Nanoparticulates As Drug Carriers

EDITOR VLADIMIR P TORCHILIN 
Northeastern University, USA 


Contents 
2 Nanoparticle Engineering 30 
2.1 Drug release mechanisms 32 
3 Site-specific Targeting with Nanoparticles: Importance of Size 
and Surface Properties 33 
4 Conclusions 37 
References 38 
4. Genetic Vaccines: A Role for Liposomes 43 
Gregory Gregoriadis, Andrew Bacon, Brenda McCormack and Peter Laing 
1 Introduction 43 
2 The DNA Vaccine 44 
3 DNA Vaccination via Liposomes 45 
3.1 Procedure for the entrapment of plasmid DNA into liposomes 46 
3.2 DNA immunization studies 47 
3.3 Induction of a cytotoxic T lymphocyte (CTL) response by 
liposome-entrapped plasmid DNA 50 
4 The Co-delivery Concept 51 
References 53 
5. Polymer Micelles as Drug Carriers 57 
Elena V. Batrakova, Tatiana K. Bronich, Joseph A. Vetro and 
Alexander V. Kabanov 
1 Introduction 57 
2 Polymer Micelle Structures 58 
2.1 Self-assembled micelles 58 
2.2 Unimolecular micelles 61 
2.3 Cross-linked micelles 62 
3 Drug Loading and Release 63 
3.1 Chemical conjugation 63 
3.2 Physical entrapment 64 
3.3 Polyionic complexation 66 
4 Pharmacokinetics and Biodistribution 68 
5 Drug Delivery Applications 72 
5.1 Chemotherapy of cancer 72 
5.2 Drug delivery to the brain 76 
5.3 Formulations of antifungal agents 77 
5.4 Delivery of imaging agents 77 
5.5 Delivery of polynucleotides 78 
Contents xv 
6 Clinical Trials 79 
7 Conclusions 79 
References 80 
6. Vesicles Prepared from Synthetic Amphiphiles — Polymeric Vesicles 
and Niosomes 95 
Ijeoma Florence Uchegbu and Andreas G. Schatzlein 
1 Introduction 95 
2 Polymeric Vesicles 96 
2.1 Polymer self assembly 97 
2.2 Polymers bearing hydrophobic pendant groups 98 
2.3 Block copolymers 101 
2.4 Preparing vesicles from self-assembling polymers 102 
2.5 Self assembling polymerizable monomers 103 
3 Polymeric Vesicle Drug Delivery Applications 104 
3.1 Drug targeting 104 
3.2 Gene delivery 105 
3.3 Responsive release 106 
3.3.1 pH 106 
3.3.2 Enzymatic 106 
3.3.3 Magnetic 107 
3.3.4 Oxygen 108 
4 Non-ionic Surfactant Vesicles (Niosomes) 108 
4.1 Self assembly 108 
4.2 Polyhedral vesicles and giant vesicles (Discomes) Ill 
4.3 Vesicle preparation 113 
5 Niosome Delivery Applications 113 
5.1 Drug targeting 113 
5.1.1 Anti cancer drugs 113 
5.1.2 Anti infectives 115 
5.1.3 Delivery to the brain 115 
5.2 Topical use of niosomes 116 
5.2.1 Transdermal 116 
5.2.2 Ocular 116 
5.3 Niosomal vaccines 116 
5.4 Niosomes as imaging agents 117 
6 Conclusions 117 
References 117 
xv i Contents 
7. Recent Advances in Microemulsions as Drug Delivery Vehicles 125 
M Jayne Lawrence and Warankanga Warisnoicharoen 
1 Definition 125 
1.1 Microemulsion versus an emulsion 125 
1.2 Microemulsion versus a nanoemulsion 126 
1.3 Microemulsions 128 
1.4 Microemulsions, swollen micelles, micelles 129 
1.5 Microemulsions and cosolvent systems 130 
2 Microemulsions as Drug Delivery Systems 130 
2.1 Self-emulsifying drug delivery systems (SEDDS) 131 
2.2 Related systems 133 
2.2.1 Microemulsion gels 133 
2.2.2 Double or multiple microemulsions 134 
2.3 Processed microemulsion formulations 134 
2.3.1 Solid state or dry emulsions 134 
3 Formulation 135 
3.1 Surfactants and cosurfactants 136 
3.2 Oils 138 
3.3 Characterization 139 
4 Routes of Administration 139 
4.1 Oral 139 
4.1.1 Proteins and peptides 140 
4.1.2 Other hydrophilic molecules 141 
4.1.3 Hydrophobic drugs 142 
4.2 Buccal 144 
4.3 Parenteral 144 
4.3.1 Long circulating microemulsions 147 
4.3.2 Targeted delivery 148 
4.4 Topical delivery 148 
4.4.1 Dermal and transdermal delivery 148 
4.5 Ophthalmic 154 
4.6 Vaginal 156 
4.7 Nasal 157 
4.8 Pulmonary 158 
4.8.1 Antibacterials 159 
5 Conclusion 160 
References 160 
Contents xvii 
8. Lipoproteins as Pharmaceutical Carriers 173 
Suwen Liu, Shining Wang and D. Robert Lu 
1 Introduction 173 
2 The Structure of Lipoproteins 174 
3 Chylomicron as Pharmaceutical Carrier 175 
4 VLDL as Pharmaceutical Carrier 176 
5 LDL as Pharmaceutical Carrier 177 
5.1 LDL as anticancer drug carriers 178 
5.2 LDL as carriers for other types of bioactive compounds . . . .179 
5.3 LDL for gene delivery 179 
6 HDL as Pharmaceutical Carriers 179 
7 Cholesterol-rich Emulsions (LDE) as Pharmaceutical Carriers . . . .180 
8 Concluding Remark 181 
References 182 
9. Solid Lipid Nanoparticles as Drug Carriers 187 
Karsten Mader 
1 Introduction: History and Concept of SLN 187 
2 Solid Lipid Nanoparticles (SLN) Ingredients and Production . . . .188 
2.1 General ingredients 188 
2.2 SLN preparation 189 
2.2.1 High shear homogenization and ultrasound 189 
2.3 High pressure homogenization (HPH) 189 
2.4 Hot homogenization 190 
2.5 Cold homogenization 190 
2.5.1 SLN prepared by solvent emulsification / 
evaporation 191 
2.5.2 SLN preparations by solvent injection 191 
2.5.3 SLN preparations by dilution of microemulsions or 
liquid crystalline phases 192 
2.6 Further processing 193 
2.6.1 Sterilization 193 
2.6.2 Drying by lyophilization, nitrogen purging and 
spray drying 194 
3 SLN Structure and Characterization 196 
4 The "Frozen Emulsion Model" and Alternative SLN Models . . . . 200 
5 Nanostructured Lipid Carriers (NLC) 201 
6 Drug Localization and Release 202 
xviii Contents 
7 Administration Routes and In Vivo Data 203 
8 Summary and Outlook 205 
References 205 
10. Lipidic Core Nanocapsules as New Drug Delivery Systems 213 
Patrick Saulnier and Jean-Pierre Benoit 
1 Introduction 213 
2 Lipidic Nanocapsule Formulation and Structure 215 
2.1 Process 215 
2.2 Influence of the medium composition 216 
2.3 Structure and purification of the LNC by dialysis 217 
2.4 Imagery techniques 218 
3 Electrical and Biological Properties 219 
3.1 Electro kinetic comportment 219 
3.2 Evaluation of complement system activation 220 
4 Pharmacokinetic Studies and Biodistribution 220 
5 Drug Encapsulation and Release 222 
5.1 Ibuprofene 222 
5.2 Amiodarone 223 
6 Conclusions 223 
References 224 
11. Lipid-Coated Submicron-Sized Particles as Drug Carriers 225 
Evan C. linger, Reena Zutshi, Terry O. Matsunaga and Rajan Ramaswami 
1 Technology 225 
2 Ultrasound Contrast Agents 228 
3 Sonothrombolysis ^r_._ 232 
4 Clinical Studies 237 
5 Blood Brain Barrier 239 
6 Drug Delivery 242 
6.1 Targeted bubbles 242 
6.2 Targeted submicron-sized droplets 244 
7 Gene Delivery 245 
8 Oxygen Delivery 247 
9 Pulmonary Delivery 248 
10 Conclusion 249 
References 250 
Contents xix 
Nanocapsules: Preparation, Characterization and Therapeutic 
Applications 255 
Ruxandra Grefand Patrick Couvreur 
1 Introduction 255 
2 Preparation 257 
2.1 Nanocapsules obtained by interfacial polymerization 257 
2.1.1 Oil-containing nanocapsules 257 
2.1.2 Nanocapsules containing an acqueous core 259 
2.2 Nanocapsules obtained from preformed polymers 261 
3 Characterization 263 
4 Drug Release 265 
5 Applications 266 
5.1 Oral route 266 
5.2 Parenteral route 267 
5.3 Ocular delivery 269 
6 Conclusion 270 
References 271 
Dendrimers as Nanoparticulate Drug Carriers 277 
Sbnke Svenson and Donald A. Tomalia 
1 Introduction 277 
2 Nanoscale Containers — Micelles, Dendritic Boxes, Dendrophanes, 
and Dendroclefts 279 
2.1 Dendritic micelles 279 
2.2 Dendritic box (Nano container) 280 
2.3 Dendrophanes and dendroclefts 282 
3 Dendrimers in Drug Delivery 282 
3.1 Cisplatin 283 
3.2 Silver salts 285 
3.3 Adriamycin, methotrexate, and 5-fluorouracil 285 
3.4 Etoposide, mefenamic acid, diclofenac, and venlafaxine . . . . 286 
3.5 Ibuprofen, indomethacin, nifedipine, naproxen, paclitaxel, 
and methylprednisolone 287 
3.6 Doxorubicin and camptothecin — self-immolative dendritic 
prodrugs 289 
3.7 Photodynamic therapy (PDT) and boron neutron capture 
therapy (BNCT) 291 
Contents 
4 Nano-Scaffolds for Targeting Ligands 292 
4.1 Folic acid 292 
4.2 Carbohydrates 293 
4.3 Antibodies and biotin-avidin binding 294 
4.4 Penicillins 295 
5 Dendrimers as Nano-Drugs 295 
6 Routes of Application 296 
7 Biocompatibility of Dendrimers 297 
8 Conclusions 299 
References 299 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly 
Soluble Drugs 307 
Rainer H. Muller and Jens-Uive A. H. Junghanns 
1 Introduction 307 
2 Definitions 308 
3 Physicochemical Properties of Drug Nanocrystals 309 
3.1 Change of dissolution velocity 309 
3.2 Saturation solubility 309 
3.3 Does size really matter? 311 
3.4 Effect of amorphous particle state 312 
4 Production Methods 313 
4.1 Precipitation methods 313 
4.1.1 Hydrosols 313 
4.1.2 Amorphous drug nanoparticles (NanoMorph®) . . . .313 
4.2 Homogenization methods 314 
4.2.1 Microfluidizer technology 314 
4.2.2 Piston-gap homogenization in water (Dissocubes®) . . 314 
4.2.3 Nanopure technology 315 
4.3 Combination Technologies 315 
4.3.1 Microprecipitation™ and High Shear Forces 
(NANOEDGE™) 315 
4.3.2 Nanopure® XP technology 316 
5 Application Routes and Final Formulations 317 
5.1 Oral administration 317 
5.2 Parenteral administration 319 
5.3 Miscellaneous administration routes 321 
6 Nanosuspensions as Intermediate Products 322 
Contents xxi 
7 Perspectives 324 
References 324 
Cells and Cell Ghosts as Drug Carriers 329 
Jose M. Lanao and M. Luisa Sayalero 
1 Introduction 329 
2 Bacterial Ghosts 329 
2.1 Application of bacterial ghosts as a delivery system 331 
3 Erythrocyte Ghosts 333 
3.1 Applications of erythrocyte ghosts as a delivery system . . . .335 
4 Stem Cells 338 
5 Polymorphonuclear Leucocytes 340 
6 Apoptopic Cells 340 
7 Tumor Cells 340 
8 Dendritic Cells 341 
9 Conclusions 341 
References 342 
Cochleates as Nanoparticular Drug Carriers 349 
Leila Zarif 
1 Introduction 349 
2 Cochleates Nanoparticles in Oral Delivery 350 
2.1 Cochleate structure 350 
2.2 Cochleate preparation 350 
2.2.1 Which phospholipid and which cation to use? 350 
2.2.2 Which molecules can be entrapped in cochleates 
nanoparticles 352 
2.2.3 Multiple ways of preparing cochleates 353 
2.3 Cochleates as oral delivery system for antifungal agent, 
amphotericin B 355 
2.3.1 In candidiasis animal model 355 
2.3.2 In aspergillosis animal model 355 
2.3.3 In cryptococcal meningitis animal model 357 
2.3.4 Toxicity of amphotericin B cochleates 357 
2.3.5 Pharmacokinetics of amphotericin B cochleates . . . . 357 
2.4 Other potential applications for cochleates 359 
2.4.1 Cochleate for the delivery of antibiotics 359 
2.4.2 Delivery of clofazimine 360 
xxii Contents 
2.4.3 Delivery of tobramycin 360 
2.4.4 Cochleate for the delivery of anti-inflammatory 
drugs 361 
2.5 Other uses of cochleates 361 
3 Conclusion 361 
References 362 
17. Aerosols as Drug Carriers 367 
N. Renee Labiris, Andrew P. Bosco and Myrna B. Dolovich 
1 Introduction 367 
2 Pulmonary Drug Delivery Devices 369 
2.1 Nebulizers 369 
2.2 Metered-dose inhalers 371 
2.3 Dry powder inhalers 373 
3 Aerosol Particle Size 373 
4 Targeting Drug Delivery in the Lung 376 
5 Clearance of Particles from the Lung 378 
5.1 Airway geometry and humidity 378 
5.2 Lung clearance mechanisms 379 
6 Nanoparticle Formulations for Inhalation 381 
6.1 Diagnostic imaging 382 
6.2 Vaccine delivery 383 
6.3 Anti Tuberculosis therapy 385 
6.4 Gene therapy 386 
7 Conclusion 388 
References 388 
18. Magnetic Nanoparticles as Drug Carriers 397 
Urs O. Hafeli and Mathieu Chastellain 
1 Introduction 397 
2 Definitions 398 
2.1 Properties of magnetic materials 398 
2.2 Nanoparticles 400 
3 Magnetic Nanoparticles 401 
3.1 Iron oxide based magnetic nanoparticles 401 
3.2 Cobalt based magnetic nanoparticles 402 
3.3 Iron based magnetic particles 402 
3.4 Encapsulated magnetic nanoparticles 403 
3.5 Biocompatibility issues of magnetic nanoparticles 403 
Contents xxiii 
4 Application of Magnetic Nanoparticles as Drug Carriers 404 
4.1 Magnetic hyperthermia 405 
4.2 Magnetic chemotherapy 406 
4.3 Other magnetic treatment approaches 408 
4.4 Magnetic gene transfer 409 
5 Conclusions 410 
References 411 
19. DQAsomes as Mitochondria-Specific Drug and DNA Carriers 419 
Volkmar Weissig 
1 Introduction 419 
2 The Self Assembly Behavior of Bis Quinolinium Derivatives 420 
2.1 Monte Carlo computer simulations 420 
2.2 Physico-chemical characterization 421 
2.3 Structure activity relationship studies 422 
3 DQAsomes as Mitochondrial Transfection Vector 424 
4 DQAsomes as Carriers of Pro-apoptotic Drugs 429 
5 Summary 432 
References 432 
20. Liposomal Drug Carriers in Cancer Therapy 437 
Alberto A. Gabizon 
1 Introduction 437 
2 The Challenge of Cancer Therapy 439 
3 The Rationale for the Use of Liposomal Drug Carriers in Cancer . . 442 
4 Liposome Formulation and Pharmacokinetics — Stealth 
Liposomes 445 
5 Preclinical Observations with Liposomal Drug Carriers 
in Tumor Models 448 
6 Liposomal Anthracyclines in the Clinic 449 
6.1 Doxil 450 
6.2 Myocet 454 
6.3 Daunoxome 454 
7 Clinical Development of Other Liposome-entrapped 
Cytotoxic Agents 455 
8 The Future of Liposomal Nanocarriers 456 
References 457 
xxiv Contents 
21. Nanoparticulate Drug Delivery to the Reticuloendothelial System 
and to Associated Disorders 463 
Mukul Kumar Basu and Sanchaita Lala 
1 Introduction 463 
2 Reticuloendothelial System and Associated Disorders 464 
3 Uptake of Nanoparticles by the Reticuloendothelial System 464 
3.1 Sites of uptake 464 
3.2 Mechanism of uptake 465 
3.3 Factors influencing uptake 468 
3.4 Role of surface modifications on uptake 469 
4 Active Targeting of Nanoparticles by Receptor Mediated 
Endocytosis 471 
5 Application in Chemotherapy 473 
6 Summary 475 
References 477 
22. Delivery of Nanoparticles to the Cardiovascular System 481 
Ban-An Khazv 
1 Introduction 481 
2 Targeting the Myocardium with Immunoliposomes 481 
3 Other Nanoparticle-Targeting of the Cardiovascular System 484 
4 Novel Application of Nano-Immunoliposomes 485 
5 CSIL as Targeted Gene or Drug Delivery 492 
6 Conclusion 495 
References 496 
23. Nanocarriers for the Vascular Delivery of Drugs to the Lungs 499 
Thomas Dziubla and Vladimir Muzykantov 
1 Introduction 500 
2 Biomedical Aspects of Drug Delivery to Pulmonary Vasculature . . 500 
2.1 Routes for pulmonary drug delivery: Intratracheal vs 
vascular 501 
2.2 Pulmonary vasculature as a target for drug delivery 501 
3 Pulmonary Targeting of Nanocarriers 503 
3.1 Effects of carrier size on circulation and tissue distribution . .503 
Contents xxv 
3.2 Passive targeting 505 
3.2.1 Mechanical retention 505 
3.2.2 Charge-mediated retention and non-viral gene 
delivery 506 
3.2.3 Pulmonary enhanced permeation-retention (EPR) 
effect 507 
3.3 Active targeting 507 
4 Carrier Design 509 
4.1 Biocompatibility 509 
4.2 Material selection (by application) 510 
4.2.1 Imaging 510 
4.2.2 Gene delivery 510 
4.2.3 Delivery of therapeutic enzymes 511 
4.2.4 Small molecule drugs 512 
4.3 Types of nanocarriers 512 
4.4 Mechanisms of drug loading 512 
4.5 Drug release mechanisms 515 
4.6 Nanocarriers for active targeting 516 
5 Conclusion: Safety Issues, Limitations and Perspectives 517 
References 518 
24. Nanoparticulate Carriers for Drug Delivery to the Brain 527 
Jorg Kreuter 
1 Introduction 527 
2 Nanoparticles 528 
3 Biodistribution 530 
3.1 Influence of surfactants on the biodistribution of 
nanoparticles 530 
3.2 Influence of PEGylation on the biodistribution of 
nanoparticles 532 
4 Pharmacology 534 
5 Brain Tumors 536 
6 Toxicology 538 
7 Mechanism of the Delivery of Drug Across the Blood-Brain 
Barrier with Nanoparticles 539 
8 Summary 541 
9 Conclusions 542 
References 542 
Contents 
Nanoparticles for Targeting Lymphatics 549 
William Phillips 
1 Introduction 549 
1.1 The lymphatic vessels 550 
1.2 Lymph nodes 551 
2 Potential for Nanoparticles for Drug Delivery to Lymphatics . . . . 553 
3 Importance of Lymph Nodes for Disease Spread and 
Potential Applications of Lymph Node Drug Delivery 554 
3.1 Cancer 554 
3.2 HIV 555 
3.3 Filaria 555 
3.4 Anthrax 556 
3.5 Tuberculosis 556 
3.6 Importance of lymph node antigen delivery for development 
of an immune response 557 
4 Factors Influencing Nanoparticle Delivery to Lymph Nodes 559 
4.1 Nanoparticle size 559 
4.2 Nanoparticle surface 559 
4.3 Effect of massage on lymphatic clearance of subcutaneously 
injected liposomes 560 
4.4 Macrophage phagocytosis 561 
4.5 Fate of nanoparticles in lymph nodes 561 
5 Nanoparticle Diagnostic Imaging Agents for Determining Cancer 
Status of Lymph Nodes 561 
5.1 Subcutaneous injection of iodinated nanoparticles for 
computed tomography imaging 561 
5.2 Subcutaneous and intraorgan injection of magnetic 
resonance (MRI) contrast agents 563 
5.3 Intravenous injection of magnetic nanoparticles for 
MRI imaging 563 
5.4 Nanoparticle diagnostic agents for localizing the sentinel 
lymph node 565 
5.5 Radiolabeled nanoparticles for sentinel lymph node 
identification 566 
5.6 99mTc-Colloidal nanoparticles for sentinel node identification . 566 
5.7 Optical 568 
5.8 Ultrasound nanobubbles 569 
6 Recently Introduced Medical Imaging Devices for Monitoring 
Lymph Node Delivery and Therapeutic Response 569 
Contents xxvii 
7 Nanoparticle Lymph Node Drug Delivery 571 
7.1 Confusion in reporting lymph node delivery 571 
7.2 Calculation of lymph node retention efficiency 573 
8 Specific Types Nanoparticles for Lymph Node Targeting 573 
8.1 PLGA nanoparticles 573 
8.2 Micelles 574 
8.3 Liposomes 574 
9 Avidin Biotin-Liposome Lymph Node Targeting Method 577 
10 Massage and the Avidin-Biotin Liposome Targeting Method 578 
11 Nanoparticles for Lymph Node Anti-Infectious Agent Delivery . . . 580 
12 Liposomes for Intraperitoneal Lymph Node Drug Delivery 581 
12.1 Intraperitoneal liposome encapsulated drugs 582 
12.2 Effect of liposome size on intraperitoneal clearance 583 
12.3 Avidin/Biotin-liposome system for intraperitoneal and 
lymph node drug delivery 584 
12.4 Mediastinal lymph node drug delivery with avidin-biotin 
system by intrapleural injection 585 
12.5 Avidin biotin for diaphragm and mediastinal lymph node 
targeting 586 
13 Nanoparticles for Cancer Therapy 587 
13.1 Intralymphatic drug delivery to lymph nodes 587 
13.2 Nanoparticles for treatment of metastatic lymph nodes of 
upper GI malignacies 589 
13.3 Lessons from endolymphatic radioisotope therapy 591 
14 Advantages of Nanoparticles for Lymphatic Radiotherapy 592 
15 Intraoperative Radiotherapy for Positive Tumor Margins 
and for Treatment of Lymph Nodes 593 
16 Potential of Using Radiolabeled Nanoparticles for Intra tumoral 
Radionuclide Therapy 593 
17 Liposome Pharmacokinetics after Intra tumoral Administration . . .595 
18 Rhenium-Labeled Liposomes for Tumor Therapy 595 
19 Nanoparticles for Immune Modulation 597 
20 Conclusions 598 
References 598 
26. Polymeric Nanoparticles for Delivery in the Gastro-Intestinal Tract 609 
Mayank D. Bhavsar, Dinesh B. Shenoy and Mansoor M. Amiji 
1 Oral Drug Delivery 609 
Contents 
2 Anatomical and Physiological Considerations of Gastro-intestinal 
Tract (GIT) for Delivery 610 
3 Introduction to Polymeric Nanoparticles as Carriers 614 
4 Preparation of Polymeric Nanoparticles 615 
5 Design Consideration for Nanoparticle-based Delivery Systems . . 619 
5.1 Polymer characteristics 619 
5.2 Drug characteristics 620 
5.3 Application characteristics 621 
6 Nanoparticles in Experimental and Clinical Medicine 621 
6.1 Drug delivery in the oral cavity 621 
6.2 Gastric mucosa as a target for oral nanoparticle-mediated 
therapy 625 
6.3 Nanoparticles for delivery of drugs and vaccines in the small 
intestine 626 
6.4 Nanoparticles for colon-specific delivery 632 
7 Integrating Polymeric Nanoparticles and Dosage Forms 634 
8 Toxicology and Regulatory Aspects 636 
8.1 Safety 637 
8.2 Quality of material/characterization 638 
8.3 Environmental considerations 638 
9 Conclusion and Outlook 639 
References 640 
Nanoparticular Carriers for Ocular Drug Delivery 649 
Alejandro Sanchez and Maria J. Alonso 
1 Biopharmaceutical Barriers in Ocular Drug Delivery. Classification 
of Nanoparticulate Carriers for Ocular Drug Delivery 650 
2 Nanoparticulate Polymer Compositions as Topical Ocular Drug 
Delivery Systems 651 
2.1 First generation: Polymer nanoparticles and nanocapsules 
for topical ocular drug delivery 652 
2.1.1 Acrylic polymers-based nanoparticles 654 
2.1.2 Polyester-based nanoparticles and nanocapsules . . .655 
2.1.3 Polysaccharide-based nanoparticles 657 
2.2 Second nanoparticles generation: The coating approach . . . . 659 
2.2.1 Polyacrylic coating 659 
2.2.2 Polysaccharide coating 660 
2.2.3 Polyethyleneglycol (PEG) coating 662 
Contents xxix 
2.3 Third nanoparticles generation: Towards functionalized 
nanocarriers 663 
3 Nanoparticulate Polymer Compositions as Subconjuctival Drug 
Delivery Systems 665 
4 Nanoparticulate Polymer Compositions as Intravitreal Drug 
Delivery Systems 665 
5 Conclusions and Outlook 667 
References 668 
Nanoparticles and Microparticles as Vaccine Adjuvants 675 
Janet R. Wendorf, Manmohan Singh and Derek T. O'Hagan 
1 Introduction 675 
2 Nanoparticle and Microparticle Preparation Methods 678 
2.1 Nanoparticles and microparticles made from polyesters . . . . 678 
2.2 Nanoparticles and microparticles made with chitosan 681 
2.3 Other nanoparticles and microparticles 681 
3 Adjuvant Effect of Nanoparticles and Microparticles 681 
3.1 Nanoparticles and microparticles as mucosal adjuvants . . . . 682 
3.2 Nanoparticles and microparticles as systemic adjuvants . . . . 686 
4 Delivery of DNA Using Nanoparticles and Microparticles 688 
5 Conclusions 690 
References 691 
Pharmaceutical Nanocarriers in Treatment and Imaging of Infection 697 
Raymond M. Schijfelers, Gert Storm and Irma A. J. M. Bakker-Woudenberg 
1 Introduction 697 
2 Carriers that are Easily Recognized as Foreign Materials 698 
3 Carriers that Avoid Recognition as Foreign Materials 701 
4 Local Application of Carriers 705 
5 Concluding Remarks 706 
References 707 
713 
1 
Introduction. Nanocarriers for Drug 
Delivery: Needs and Requirements 
Vladimir Torchilin 
Fast developing nanotechnology, among other areas, is expected to have a dramatic 
impact on medicine. The application of nanotechnology for treatment, diagnosis, 
monitoring, and control of biological systems has recently been determined by the 
NIH as nanomedicine. Among the approaches for exploiting nanotechnology developments 
in medicine, various nanoparticulates offer some unique advantages as 
pharmaceutical delivery systems and image enhancement agents.1,2 Several varieties 
of nanoparticles are available3: different polymeric and metal nanoparticles, 
liposomes, micelles, quantum dots, dendrimers, microcapsules, cells, cell ghosts, 
lipoproteins, and many different nanoassemblies. All of these nanoparticles can 
play a major role in diagnosis and therapy. This book is attempting to present the 
broad overview of different nanoparticulate drug delivery systems with all their 
advantages and limitations, as well as potential areas of their clinical applications. 
The paradigm of using nanoparticulate pharmaceutical carriers to enhance the 
in vivo efficiency of many drugs, anti-cancer drugs, first of all, well established 
itself over the past decade both in pharmaceutical research and clinical setting, and 
does not need any additional proofs. Numerous nanoparticle-based drug delivery 
and drug targeting systems are currently developed or under development.4,5 
Their use aims to minimize drug degradation upon administration, prevent undesirable 
side effects, and increase drug bioavailability and the fraction of the drug 
accumulated in the pathological area. Pharmaceutical drug carriers, especially the 
1 
2 Torchilin 
ones for parenteral administration, are expected to be easy and reasonably cheap 
to prepare, biodegradable, have small particle size, possess high loading capacity, 
demonstrate prolonged circulation, and, ideally, specifically or non-specifically 
accumulate in required pathological sites in the body.6 
High molecular weight (40 kDa or higher), long-circulating macromolecules, 
including proteins and peptides, conjugated with water-soluble polymers, are capable 
of spontaneous accumulations in various pathological sites such as solid tumors, 
infarcts, and inflammations via the enhanced permeability and retention effect 
(EPR).7'8 This effect is based on the fact that pathological (tumor, infarct) vasculature, 
unlike vasculature of healthy tissues, is "leaky", i.e. penetrable for macromolecules 
and nanoparticulates which allows for macromolecules to accumulate 
in the pathological tissue (such as interstitial tumor space). In the case of tumors, 
such accumulation is also facilitated by the fact that lymphatic system, responsible 
for the drainage of macromolecules from normal tissues, is virtually not working 
in case of many tumors as the result of the disease.8 It has been found that the 
effective pore size of most peripheral human tumors range from 200 nm to 600 nm 
in diameter, with a mean of about 400 nm. The EPR effect allows for passive targeting 
to tumors and other pathological sites based on the cut-off size of the leaky 
vasculature.9 
Among particulate drug carriers, liposomes, micelles and polymeric nanoparticles 
are the most extensively studied and possess the most suitable characteristics 
for encapsulation of many drugs and diagnostic (imaging) agents. Many other systems 
meeting certain more specific requirements (and reviewed in this book) are 
also suggested and currently under development. Making these nanocarriers multifunctional 
and stimuli-responsive can dramatically enhance the efficiency of various 
drugs carried by these carriers. These functionalities are expected to provide: 
(a) prolonged circulation in the blood10'11 and the ability to accumulate in various 
pathological areas (such as solid tumors) via the EPR effect (protective polymeric 
coating with PEG is used for this purpose)12,13; (b) ability to specifically recognize 
and bind target tissues or cells via the surface-attached specific ligand (monoclonal 
antibodies as well as their Fab fragments and some other molecules are used for this 
purpose)14; (c) ability to respond local stimuli characteristic of the pathological site 
by, for example, releasing an entrapped drug or specifically acting on cellular membranes 
under the abnormal pH or temperature in disease sites (this property could 
be provided by surface-attached pH- or temperature-sensitive coatings); (d) ability 
to penetrate inside cells bypassing the lysosomal degradation for efficient targeting 
of intracellular drug targets (for this purpose, the surface of nanocarriers may be 
decorated by cell-penetrating peptides). Those are just the most evident examples. 
Some other specific properties can also be listed, such as an attachment of diagnostic 
moieties. Even the use of individual functionalities is already associated with highly 
Introduction. Nanocarriers for Drug Delivery: Needs and Requirements 3 
positive clinical outcome — the success of Doxil®, doxorubicin in long-circulating 
PEG-coated liposome, represents a good example.15 
In addition, there are numerous engineered constructs, assemblies, architectures, 
and particulate systems, whose unifying feature is the nanometer scale size 
range (from a few to 250 nm). Together with already listed systems, these include 
cyclodextrins, niosomes, emulsion particles, solid lipid particles, drug nanocrystals, 
metal and ceramic nanoparticles, protein cage architectures, viral-derived capsid 
nanoparticles, polyplexes, cochleates, and microbubbles.4,5,16-19 Therapeutic and 
diagnostic agents can be encapsulated, covalently attached, or adsorbed on to such 
nanocarriers. These approaches can easily overcome drug solubility issues, particularly 
with the view that large proportions of new drug candidates emerging from 
high-throughput drug screening initiatives are water insoluble. Yet, some carriers 
have a low capacity to incorporate active compounds (e.g. dendrimers, whose size 
is in the order of 5-10 nm). There are alternative nanoscale approaches for solubilization 
of water insoluble drugs too.20-23 One approach is to mill the substance 
and then stabilize smaller particles with a coating; this forms nanocrystals in size 
ranges suitable for oral delivery, as well as for intravenous injection.24,25 Pharmacokinetic 
profiles of injectable nanocrystals may vary from rapidly soluble to slowly 
dissolving in the blood. 
In general, the development of drug nanocarriers for poorly soluble pharmaceuticals 
represents a special task and still faces some unresolved issues. The therapeutic 
application of hydrophobic, poorly water-soluble agents is associated with 
some serious problems, since low water-solubility results in poor absorption and 
low bioavailability.26 In addition, drug aggregation upon intravenous administration 
of poorly soluble drugs might lead to such complications as embolism27 and 
local toxicity.28 On the other hand, the hydrophobicity and low solubility in water 
appear to be intrinsic properties of many drugs,29 since it helps a drug molecule to 
penetrate a cell membrane and reach important intracellular targets.30,31 To overcome 
the poor solubility of certain drugs, clinically acceptable organic solvents are 
used in their formulations,28 as well as liposomes32 and cyclodextrins.16 Another 
alternative is associated with the use of various micelle-forming surfactants in formulations 
of insoluble drugs. 
By virtue of their small size and by functionalizing their surface with synthetic 
polymers and appropriate ligands, nanoparticulate carriers can be targeted 
to specific cells and locations within the body after intravenous and subcutaneous 
routes of injection. Such approaches may enhance detection sensitivity in medical 
imaging, improve therapeutic effectiveness, and decrease side effects. Some of the 
carriers can be engineered in such a way that they can be activated by changes in the 
environmental pH, chemical stimuli, by the application of a rapidly oscillating magnetic 
field, or by application of an external heat source.19,33-35 Such modifications 
4 Torchilin 
offer control over particle integrity, drug delivery rates, and the location of drug 
release, for example, within specific organelles. Some are being designed with the 
focus on multifunctionality; these carriers target cell receptors and delivers drugs 
and biological sensors simultaneously. Some include the incorporation of one or 
more nanosystems within other carriers, as in the micellar encapsulation of quantum 
dots; this delineates their inherent nonspecific adsorption and aggregation in 
biological environments.36 
The use of nanoparticulate drug carriers seems to be especially important 
for developing efficient anticancer therapies. Although significant advances have 
occurred in our understanding of tumor origin, growth, metastasis, and many different 
types of pharmacological agents have been developed over the years to treat 
tumors, the problem of optimum delivery remains a formidable challenge. For any 
of the drug therapy strategies to be effective, the agent must be able to reach the 
tumor mass in sufficient concentration, traverse through the tumor microcirculation, 
diffuse into the interstitium, and remain at the site for the duration to induce 
tumoricidal effect. As was already mentioned, due to the porosity of the tumor 
vasculature and the lack of lymphatic drainage, blood-borne macromolecules and 
nanoparticles are preferentially distributed in the tumor via the EPR effect. However, 
nanoparticles can also be actively targeted to tumors by modifying their 
surface with certain cell-specific ligands for receptor-mediated uptake. The use 
of specific "vector" molecules can further enhance tumor targeting of nanocarries 
or make them the EPR-effect independent. The latter is especially important 
for the cases of tumors with immature vasculature, such as tumors in the early 
stages of their development, and for delocalized tumors. Vector molecules (those 
having affinity toward ligands characteristic for target tissues), capable of recognizing 
tumors were found among antibodies, peptides, lectins, saccharides, hormones, 
transferrin and some low molecular weight compounds (riboflavin, folate). 
From this list, antibodies and their fragments provide the most universal opportunity 
to target various for targeting and have the highest potential specificity. 
Vector molecules can be used for the targeting of nanoreservoir delivery systems 
as well. PEG-modified long-circulating doxorubicin-containing immunoliposomes 
targeted with anti-HER-2/neu monoclonal antibody fragments represent a recent 
example of increased efficiency of targeted delivery systems.37 In all studied HER2- 
overexpressing models, immunoliposomes showed potent anticancer activity superior 
to that of control non-targeted liposomes. In part, this superior activity was 
attributed to the ability of the immunoliposomes to deliver their load inside the 
target cells via the receptor-mediated endocytosis, which is obviously important if 
the drug's site of action sites locates inside the cell. 
An important problem is associated with the clearance of drug carriers from the 
circulation. Nanoparticular pharmaceutical carriers administered into the systemic 
Introduction. Nanocarriers for Drug Delivery: Needs and Requirements 5 
circulation will be essentially removed within an hour of administration by the 
macrophages of the reticulo-endothelial system. To prolong the circulation of 
nanoparticles by evading the macrophages, their surface is modified with watersoluble 
polymers. Poly(ethylene glycol) (PEG) is very popular for surface modification 
of nanoparticulate drug delivery systems, since it has a long history of 
safe usage in biological and pharmaceutical products. Surface-bound PEG chains 
extend into the aqueous physiological environment and repel proteins, decrease 
antibody formation, and increase the circulation of the formulation in the plasma 
for extended period of time by the steric repulsion mechanism.38 
With rapid advances in molecular biology and genetic engineering, there is 
an unprecedented opportunity for delivery of drugs and genes to intracellular 
targets.39 In the case of cancer, for instance, the effectiveness of many anticancer 
drugs is limited due to its inability to reach the target site in sufficient concentrations 
and to exert the pharmacological effect. Current gene delivery systems are 
classified as being either viral or non-viral in origin. Viruses are efficient in delivery 
of genes; however, they suffer from poor safety profile. Non-viral gene delivery systems, 
albeit not as efficient as viruses, have promise of safety and reproducibility in 
manufacturing. To enhance delivery of drugs to intracellular targets and gene transfection 
efficiency using non-viral delivery systems, it is necessary to identify ways 
of overcoming the cellular barriers, for example, by using various cell-penetrating 
proteins and peptides.40,41 
Self-assembled nanosystems (nanoassemblies) for targeting subcellular 
organelles, such as the mitochondria, are also developed.42 It has become increasingly 
evident that mitochondrial dysfunction contributes to a variety of human 
disorders. Moreover, since the middle of 1990s, mitochondria, the "power houses" 
of the cell, have also become accepted as the cell's "arsenals", which reflects their 
increasingly acknowledged key role during apoptosis. Based on these recent developments 
in mitochondrial research, increased pharmacological and pharmaceutical 
efforts have led to the emergence of "Mitochondrial Medicine" as a whole new field 
of biomedical research. 
Nanoparticulate drug delivery systems are very important for the delivery of 
peptide and protein drugs and may represent a valid alternative to soluble polymeric 
carriers used earlier. The use of this type of carriers allows achieving much 
higher active moiety/carrier material ratio compared with "direct" molecular conjugates. 
They also provide better protection of protein and peptide drugs against 
enzymatic degradation and other destructive factors upon parenteral administration, 
because the carrier wall completely isolates drug molecules from the environment. 
All nanoparticulate carriers have the size, which excludes a possibility of renal 
filtration. Among particulate drug carriers, liposomes are the most extensively studied 
and poses the most suitable characteristics for protein (peptide) encapsulation. 
6 Torch i I in 
Similar to macromolecules, protein and peptide drug-bearing liposomes are capable 
of accumulating in tumors of various origins via the EPR effect.6-8 In some 
cases, however, the liposome size is too large to provide an efficient accumulation 
via the EPR effect presumably due to relatively small tumor vasculature cut 
off size.43,44 In such cases, alternative delivery systems with smaller sizes, such as 
micelles (prepared, for example, from PEG-phospholipid conjugates) can be used. 
These particles lack the internal aqueous space and are smaller than liposomes. 
Protein or peptide pharmaceutical agent can be covalently attached to the surface 
of these particles or incorporated into them via chemically attached hydrophobic 
group ("anchor"). 
In conclusion, even a brief listing of some key problems of nanocarrier-mediated 
drug delivery shows how broad and intense this area is. In addition to this, 
nanoscale-based delivery strategies are beginning to make a significant impact on 
global pharmaceutical planning and marketing. The leading experts in the area of 
nanparticulate-mediated drug delivery attempted to address these and many other 
topics in this book. We strongly believe that every reader will find the book useful 
and stimulating. 
References 
1. West JL and Halas NJ (2000) Applications of nanotechnology to biotechnology commentary. 
Curr Opin Biotechnol 11:215. 
2. La Van DA, Lynn DM and Langer R (2002) Moving smaller in drug discovery and 
delivery. Nat Rev Drug Discov 1:77. 
3. Sahoo SK and Labhasetwar V (2003) Nanotech approaches to drug delivery and imaging. 
Drug Discov Today 8:1112. 
4. Miiller, RH (1991) Colloidal Carriers for Controlled Drug Delivery and Targeting. 
Wissenschaftliche Verlagsgesellschaft: Stuttgart, Germany and CRC Press: Boca Raton. 
5. Cohen S and Bernstein H (eds.) (1996). Microparticulate Systems for the Delivery of Proteins 
and Vaccines. Marcel Dekker, New York. 
6. Gref R, Minamitake Y, Peracchia MT, Trubetskoy V, Torchilin VP and Langer R (1994) 
Biodegradable long-circulating polymeric nanospheres. Science 263:1600. 
7. Maeda H (2001) SMANCS and polymer-conjugated macromolecular drugs: Advantages 
in cancer chemotherapy. Adv Drug Deliv Rev 46:169. 
8. Maeda H, Sawa T and Konno T (2001) Mechanism of tumor-targeted delivery of macromolecular 
drugs, including the EPR effect in solid tumor and clinical overview of the 
prototype polymeric drug SMANCS. / Control Rel 74:47. 
9. Yuan F, Dellian M, Fukumura M, Leunig M, Berk BD, Torchilin VP and Jain RK (1995) 
Vascular permeability in a human tumor xenograft, Molecular size dependence and 
cutoff size. Cancer Res 55:3752. 
10. Lasic DD and Martin F (eds.) (1995) Stealth Liposomes. CRC Press: Boca Raton. 
Introduction. Nanocarriers for Drug Delivery: Needs and Requirements 7 
11. Torchilin VP and Trubetskoy VS (1995) Which polymers can make nanoparticulate drug 
carriers long-circulating? Adv Drug Del Rev 16:141. 
12. Lukyanov AN, Hartner WC and Torchilin VP (2004) Increased accumulation of PEGPE 
micelles in the area of experimental myocardial infarction in rabbits. / Control Rel 8, 
94:187. 
13. Maeda H, Wu J, Sawa T, Matsumura Y and Hori K (2001) Tumor vascular permeability 
and the EPR effect in macromolecular therapeutics: A review. / Control Rel 65:271. 
14. Torchilin VP (1998) Polymer-coated long-circulating microparticular pharmaceuticals. 
] Microencapsulation 15:1. 
15. O'Shaughnessy JA (2003) Pegylated liposomal doxorubicin in the treatment of breast 
cancer. Clin Breast Cancer 4,318. 
16. Thompson D and Chaubal MV (2000) Cyclodextrins (CDS) — excipients by definition, 
drug delivery systems by function (Part I: Injectable applications). Drug Del Technol 2:34. 
17. Zhang L and Eisenberg A (1995) Multiple morphologies of "crew-cut" aggregates of 
polystyrene-b-poly(acrylic acid) block copolymers. Science 268:1728. 
18. Gref R, Domb A, Quellec P, Blunk T, Muller RH, Verbavatz JM and Langer R (1995) 
The controlled intravenous delivery of drugs using PEG-coated sterically stabilized 
nanospheres. Adv Drug Del Rev 16:215. 
19. Cammas S, Suzuki K, Sone C, Sakurai Y, Kataoka K and Okano T (1997) Thermorespensive 
polymer nanoparticles with a core-shell micelle structure as site specific drug 
carriers. / Control Rel 48:157. 
20. Kabanov AV, Batrakova EV and Alakhov VY (2002) Pluronic block copolymers as novel 
polymer therapeutics for drug and gene delivery. / Control Rel 82:189. 
21. Kwon GS (2003) Polymeric micelles for delivery of poorly water-soluble compounds. 
Crit Rev Ther Drug Can Syst 20:357. 
22. Jones M and Leroux J (1999) Polymeric micelles — a new generation of colloidal drug 
carriers. Eur J Pharm Biopharm 48:101. 
23. Torchilin VP (2001) Structure and design of polymeric surfactant-based drug delivery 
systems. / Control Rel 73:137. 
24. Muller RH and Keck CM (2004) Challenges and solutions for the delivery of biotech 
drugs — a review of drug nanocrystal technology and lipid nanoparticles. / Biotechnol 
113:151. 
25. Kraft WK, Steiger B, Beussink D, Quiring JN, Fitzgerald N, Greenberg HE and 
Waldman SA (2004) The pharmacokinetics of nebulized nanocrystal budesonide suspension 
in healthy volunteers. / Clin Pharmacol 44:67. 
26. Lipinski CA, Lombardo F, Dominy BW and Feeney PJ (2001) Experimental and computational 
approaches to estimate solubility and permeability in drug discovery and 
development settings. Adv Drug Del Rev 46:3. 
27. Fernandez AM, Van Derpoorten K, Dasnois L, Lebtahi K, Dubois V, Lobl TJ, Gangwar S, 
Oliyai C, Lewis ER, Shochat D and Trouet A (2001) N-Succinyl-(beta-alanyl-L-leucyl- 
L-alanyl-L-leucyl) doxorubicin: An extracellularly tumor-activated prodrug devoid of 
intravenous acute toxicity. / Med Chem 44:3750. 
8 Torchilin 
28. Yalkowsky SH (ed.) (1981) Techniques of Solubilization of Drugs. Marcel Dekker: New York 
and Basel. 
29. Shabner BA and Collings JM (eds.) (1990) Cancer Chemotherapy: Principles and Practice. 
J. B. Lippincott Co: Philadelphia. 
30. Yokogawa K, Nakashima E, Ishizaki J, Maeda H, Nagano T and Ichimura F (1990) Relationships 
in the structure-tissue distribution of basic drugs in the rabbit. Pharm Res 
7:691. 
31. Hageluken A, Grunbaum L, Nurnberg B, Harhammer R, Schunack W and Seifert R 
(1994) Lipophilic beta-adrenoceptor antagonists and local anesthetics are effective direct 
activators of G-proteins. Biochem Pharmacol 47:1789. 
32. Lasic DD and Papahadjopoulos (eds.) (1998) Medical Applications of Liposomes. Elsevier: 
New York. 
33. Le Garrec D, Taillefer J, VanLier JE, Lenaerts V and Leroux JC (2002) Optimizing 
pH-responsive polymeric micelles for drug delivery in a cancer photodynamic therapy 
model. / Drug Targ 10:429. 
34. Meyer O, Papahadjopoulos D and Leroux JC (1998) Copolymers of N-isopropylacrylamide 
can trigger pH sensitivity to stable liposomes. FEBS Lett 41:61. 
35. Chung JE, Yokoyama M, Yamato M, Aoyagi T, Sakurai Y and Okano T (1999) Thermoresponsive 
drug delivery from polymeric micelles constructed using block copolymers 
of poly(N-isopropylacrylamide) and poly(butylmethacrylate). / Control Rel 62:115. 
36. Stroh M, Zimmer JP, Duda DG, Levchenko TS, Cohen KS, Brown EB, Scadden DT, 
Torchilin VP, Bawendi MG, Fukumura D and Jain RK (2005) Quantum dots spectrally 
distinguish multiple species within the tumor milieu in vivo. Nat Med 11:678. 
37. Park JW, Kirpotin DB, Hong K, Shalaby R, Shao Y, Nielsen UB, Marks JD, 
Papahadjopoulos D and Benz CC (2001) Tumor targeting using anti-her2 immunoliposomes. 
/ Control Rel 74:95. 
38. Veronese FM and Harris JM (2002) Introduction and overview of peptide and protein 
pegylation. Adv Drug Del Rev 54:453. 
39. Torchilin VP and Lukyanov AN (2003) Peptide and protein drug delivery to and into 
tumors: Challenges and solutions. Drug Discov Today 8:259. 
40. Schwarze SR, Ho A, Vocero-Akbani A and Dowdy SF (1999) In vivo protein transduction: 
Delivery of a biologically active protein into the mouse. Science 285:1569. 
41. Gupta B, Levchenko TS and Torchilin VP (2005) VP: Intracellular delivery of large 
molecules and small particles by cell-penetrating proteins and peptides. Adv Drug Del 
Rev 57:637. 
42. Weissig V (2003) Mitochondrial-targeted drug and DNA delivery. Crit Rev Ther Drug 
CarrSyst 20:1. 
43. Weissig V, Whiteman KR and Torchilin VP (1998) Accumulation of protein-loaded longcirculating 
micelles and liposomes in subcutaneous Lewis lung carcinoma in mice. Pharm 
Res 15:1552. 
44. Hobbs SK, Monsky WL, Yuan F, Roberts WG, Griffith L, Torchilin VP and Jain RK (1998) 
Regulation of transport pathways in tumor vessels: Role of tumor type and microenvironment. 
Proc Natl Acad Sci USA 95:4607. 
2 
Nanoparticle Flow: Implications 
for Drug Delivery 
Alexander T. Florence 
1. Introduction 
While the experimental study of nanoparticle flow in vivo proves to be difficult, a 
variety of theoretical and practical techniques are becoming available to allow some 
understanding of the phenomena involved. These processes include (a) convective 
flow induced by the flow of blood, lymph or interstitial fluid, (b) the influence of the 
interaction of nanoparticles with themselves or with biological components and the 
effect of this on their transport, and (c) the effect of fluid flow and hence shear forces 
on particle access to, interaction with and removal from receptors. Diffusion and 
movement of particle suspensions in complex media such as interstitial tissue must 
also be considered. Much of the theoretical work which is relevant to this exploration 
of nanoparticle flow has not been directed towards biological endpoints, but 
this body of knowledge, and the analogous literature on the dynamic behavior of 
bacteria, erythrocytes and platelets provides the basis of a more rigorous analysis 
of the factors involved in drug carrier nanoparticle flow. 
As discussed in this book, nanoparticles are of value in drug, vaccine and gene 
delivery because their small dimension compared with microparticles allows them 
to interact more effectively with cells, be safely injected, and amongst other characteristics, 
diffuse further into tissues, and into and through individual cells. The 
flow of nanoparticles in capillaries, lymphatics, tumor vessels, their extravasation1 
and movement in the cytoplasm of cells are all aspects of the topic covered in 
9 
10 Florence 
this overview, albeit from a phenomenological viewpoint. It is clear that particle 
diameter is a key parameter in the characterization and behavior of nanoparticle 
suspensions. In several of the analyses here, it becomes clear that another advantage 
of nanoparticles may be the relative lack of effect of shear stress, once particles 
adhere to surfaces as a prelude to uptake; this is contrasted with targeted microspheres 
whose residence on receptors and surfaces is size dependent, the larger 
particles being more vulnerable to detachment by shear forces. 
This chapter considers questions relating to the flow of nanoparticles in vivo, 
but which has often been simulated in vitro by chemical engineers and physicists 
interested in particle behavior in flow conditions. Spherical particles are the norm, 
but not all nanosystems are spherical. The influence of asymmetry on the transport 
of nanoparticles in vivo is largely unknown, although the rheological characteristics 
of asymmetric particle suspensions have been understood for a long time. Flow 
behavior of nanoparticles in complex networks of narrow capillaries has relevance 
in the design and operation of microfluidic devices, as well as in drug delivery 
and targeting, and in toxicology2; the extent to which it is relevant for delivery and 
targeting is explored here. Figure 1 illustrates diagrammatically some of the areas 
of interest. 
In physical terms, the following situations could be considered: (i) particle flow 
in rapidly flowing blood, including segregation and deposition of particles, and the 
behavior of particles at bifurcations in the capillary supply; (ii) the effect of shear 
on adhesion of particles of different size and shape; (iii) particle flow in more static 
conditions, for example, in the tumor interstitium or the lymphatics; (iv) flow of 
particles in tissues, including the flow of particles into narrow pores and (v) flow or 
diffusion within the anisotropic interior of cells. Added complications arise where 
bioadhesive or ligand-decorated particles are involved. In the latter case, binding 
and flow are linked. 
The enquiry can be divided into a discussion of the flow and movement of 
nanoparticles as a function of their size and route of administration, the influence 
of convection and blood or lymph flow, the influence of flow dynamics on the 
interaction of nanoparticles with target tissues and receptors, and the movement 
of nanoparticles once they have been absorbed and are making their way through 
individual cells and tissues towards a target. 
The routes of administration where flow is important potentially include all, 
even the oral route where particle flow and dynamics in the gut lumen and in the 
vicinity of both villi and microvilli is important. Particles below a critical size are 
taken up by the M-cells of Peyer's patches and by normal enterocytes, albeit in 
small quantities and find themselves in the lymph vessels, lymph nodes, blood, 
liver and spleen.3'4 If the flow of nanoparticles away from their site of absorption 
is restricted due to the flow in lymph or blood being slow, this will reduce 
Nanoparticle Flow: Implications for Drug Delivery 11 
* II 
Fig. 1. A diagrammatic representation (not to scale) of some of the areas where flow and 
transport of nanoparticles is key. I: flow in the GI tract after oral administration; II: access to and 
adhesion to M-cells of Peyer 's patches or to enterocy tes; III: passage into the mesenteric lymph; 
IV: flow in the lymph vessels and entrapment in the lymph nodes (not shown); V: transport 
between lymph and blood. A: blood flow; B: adhesion to capillary walls; C: extravasation and 
flow in tissue; D: flow and deposition at vessel bifurcations; and E: movement into tumours. 
Each route (the subcutaneous route is also indicated) will involve a complex sequence of 
nanoparticle pathways, most involving lymph, blood and intestinal fluid. 
bioavailability and distribution. Rapid flow provides superior sink conditions and 
hence the use of everted gut sacs and cell monolayers in vitro can give unrealistic 
results for nanoparticle transit. The influence of flow dynamics on extravasation and 
perhaps on the enhanced permeability and retention (EPR) effect for nanocarriers 
has perhaps not been fully addressed. Both involve consideration of particle size, the 
diffusion and flow of nanoparticles through narrow channels, as well as navigation 
of tortuous environments. The availability of "extreme" nanoparticles in the form 
of dendrimers5 and quantum dots6 makes this topic a vital one in understanding 
the fate, toxicity7 or accumulation of what is metrically a wide range of systems. 
2. Background 
Our own interest in this field has resulted in part from studies on the size dependency 
of nanoparticle uptake after oral administration, where mesenteric lymphatic 
transport of 500 nm nanoparticles post absorption is determined by the flow of particles 
in a single file in the smallest mesenteric lymph vessels (Fig. 2). In addition, 
12 Florence 
Fig. 2. 500 nm polystyrene latex particles flowing in the mesenteric lymph vessels mostly 
in single-file mode, from Jani et al. 
studies on the flow of liposomes and niosomes, and the extrusion of flexible vesicles 
from glass capillaries under pressure, converting polyhedral vesicles into multilayer 
tubules, led us to consider the influence of stress forces on carrier integrity. 
Clearly, the elasticity of vesicles is important in their negotiation of capillaries when 
their diameter exceeds that of the capillary. The flow of particles in fabricated capillaries 
which have a radius close to the particle radius is a challenge that has been 
tackled theoretically.8'1* We have suggested that multi-bilayer tubules (Fig. 3) can 
act as models for such flow experiments.11 
Rheological examination of nanoparticle-blood mixtures and nanoparticle suspensions 
of mixed radius has also illustrated the potential complexity of particle 
Fig. 3. A flexible non-ionic surfactant based multi-bilayer tube, around 1 /xm in diameter, 
extruded from a suspension of polyhedral niosomes, which might be adapted for use as a 
model for the study of capillary nanoparticle flow.11 
Nanoparticle Flow: Implications for Drug Delivery 13 
movement in blood (unpublished data). In addition, erythrocyte blockage at bifurcations, 
or narrowing of vessels can lead to slowing down of blood flow and a 
change in rheology as the haematocrit increases. More recently, investigation of 
the transport of nanoparticles across cell monolayers12 and intracellular transport 
of dendrimers13 has assisted in defining some of the issues involved in targeting 
within cells. 
3. Studies on Nanoparticle Flow 
The work of Fokin and colleagues14 on the transport of viral-sized colloids, following 
intravenous or intra-lymphatic injection, is relevant to drug delivery even if 
their objectives were different. 100-200 nm diameter sulphur colloid particles reach 
the lymph after IV injection in around 25 minutes; after intra-lymphatic injection 
particles appear in the venous blood only after 4 seconds. Following subcutaneous 
injection, similar particles14-16 reach the lymph after 2-9 min, although 95% of the 
particles remain at the injection site for at least 45 minutes. Here, the nanoparticles 
are being used as indicators of blood and lymph flow. What is also relevant to drug 
delivery is the influence of fluid flow on the movement and fate of nanoparticles. 
Ilium et al.17 observed uptake rates of 1.27 jxm and 15.8 /zm polystyrene particles 
in the lung and liver after IV injection. The sequestration in the lung was size 
dependent, but possibly affected because the smaller particles were taken up by 
the Kupffer cells of the liver, leaving the larger particles free to be taken up by the 
lung tissue. The rapidity of this suggests that, in effect, flow of the microparticles 
is solely determined by blood flow. 
4. Convection and Diffusion 
Blood flow drives the convective flow of suspended particles. Diffusional transport 
occurs in static conditions or conditions of low fluid velocity. In a tube of flowing 
liquid, convective dynamics propel the particles in the direction of flow, but at the 
walls of the tube, there is the possibility of particle diffusion resulting in deposition. 
Blood velocity (mm s_1) in arterioles and venules is a function of vessel diameter, 
as shown in Fig. 4. In venules, the maximum velocity according to Jain18 is approximately 
12 mm s_1, while in arterioles, it can reach about 30 mm s^1. 
Fluid velocity in tubes is not constant throughout the diameter of the tube as 
Fig. 5 indicates, a feature that is important when the interaction of nanoparticles 
with epithelial cells or capillary walls is considered. 
The radial variation of shear is a factor that must be considered in polydisperse 
nanoparticulate systems and where nanoparticles adhere to erythrocytes, causing 
two distinct size distributions. If nanoparticles adhere to erythrocytes20 or other 
14 Florence 
Arterioles 
o 0 o 
o 
°° s 
o 
I" I 
0 25 
Vmax(mm/s) 
o 
s 
o 
© 
r-35 
Venules 
- 30 
-25 
" 20 
o 
5 g »|5i d^^OD0 a 
) 25 50 
Vessel Diameter (nm) 
Fig. 4. Maximum blood velocity (mms 1) in arterioles and venules as a function of vessel 
diameter, redrawn from R. K. Jain.18 
Fig. 5. Diagram showing the velocity pattern in a tube of flowing liquid. Particles of different 
size separate according to their diameter. The large particles, being unable to approach 
close to the capillary wall, experience the faster fluid streamlines toward the centre; hence, 
they move more rapidly, as described by Silebi and DosRamos.18,19 This is the basis of the 
field flow fractionation. 
blood elements, the translocation of the particles is controlled by the particular 
element to which it adheres. The rheology of suspensions of mixed particles is 
complex: viscosity reduces first with an increase in the fraction of larger particles 
in a suspension, and as the volume fraction increases, so does the viscosity.21 
Ding et al.22 formulated a theoretical model examining particle migration 
in nanoparticle suspensions flowing through a pipe. "The model considers particle 
migration due to spatial gradients in viscosity and shear rate as well as 
Brownian motion. Particle migration due to these effects can result in significant 
non-uniformity in particle concentration over the cross section of the pipe" in 
particular for larger particles. Three mechanisms were proposed by the authors 
for migration in such non-uniform shear flow: (i) shear induced migration where 
Nanoparticle Flow: Implications for Drug Delivery 15 
particles move from regions of higher shear rate to regions of lower shear rate; 
(ii) viscosity gradient induced migration — particles move from regions of higher 
viscosity to regions of lower viscosity and (iii) self-diffusion due to Brownian 
motion. 
Diffusion inside microtubules has been studied to understand taxol binding to 
tubulin structures.23 The dimension of the tubulin lumen is of the order of 17nm, 
approaching macromolecular dimensions, leading to friction between the inner 
walls and the moving macromolecules. This "hindrance" will also be an issue in 
the movement of nanoparticles in the smallest capillaries. With dendrimers whose 
diameters may be as small as 6 nm, the application of hindered theory to their movement 
could be relevant. No vessels are of this small radius, but the key parameter 
is the ratio of particle diameter to capillary diameter. The approach may well be 
important in cellular networks. It is not only capillary vessels that are the conduits 
of particle movement, but after extravasation, there is passage through cellular networks. 
The process could be considered to be akin to diffusion in porous networks. 
Binding of the moving particle (or macromolecule) to the luminal surface of the vessel 
will also hinder free flow or movement, a positive event in the case of specific 
ligand targeting of "decorated" systems. 
Polydisperse nanosystems can segregate during flow or migrate differentially 
leading to concentration differences.22 Particle-image velocimetry (PIV)24 has been 
used to track the flow characteristics of microparticles. The effects of flow on adhesion 
of monocytes to endothelial cells25 is relevant to the influence of flow and shear 
in particle interactions and uptake. 
The significance of flow can be demonstrated by the use of pharmacological 
agents which change normal vessel patency, so that by the concomitant use of 
noradrenalin26 or angiotensin26,27 which constrict only normal vessels, the ratio of 
tumor to normal tissue blood flow can be optimized. 
5. Bifurcations 
Many theoretical studies of nanoparticle flow deal with linear tubes, whereas 
in vivo movement occurs through complex vessel architectures with bifurcations28,29 
(Fig. 6). Behavior at bifurcations in a vascular or capillary system is dependent not 
only on particle diameter, but also on the rigidity or flexibility of the particle concerned. 
Colloid transport in a bifurcating structure has been the subject of one recent 
paper.30 It is a process which depends on the orientation of the bifurcations, especially 
with particles whose density is greater than that of the medium, as well as on 
the different flow rates in the individual branches which are likely to be of different 
sizes. 
If nanoparticles are trapped or associate at bifurcations or indeed other obstacles 
in capillaries, then it is likely that they might associate more permanently, thereby 
16 Florence 
A •& - 
Q„.30IAnin HB M ,.., 2.»..«s.» M ae'0.3 
Fig. 6. Three-dimensional distributions of nanoparticles in a bifurcation airway model of 
Zhang et ah, Aerosol Sci., 2005, 36, 211-233. DEF is the deposition enhancement factor, the 
representations shown here being for a steady inhalation. While these data are for air-flow, 
not dissimilar patterns of deposition might be estimated to occur in liquid flows. Deposition 
in these models occurs primarily by Brownian diffusion; deposition efficiencies increase with 
decreasing nanoparticle size and lower inlet Reynolds numbers. 
changing their intrinsic rheological behavior. Flexible particles do not of course 
suffer the same constraints in movement and progress, but their flexibility can lead 
to slow negotiation of movement around obstacles (Fig. 7). 
6. Interaction with Blood Constituents and 
Endogenous Molecules 
Nanoparticles may interact with blood constituents31: the adsorption of albumin, 
IgG and fibrinogen from blood onto hydrophobic particles is well known, but the 
Nanoparticle Flow: Implications for Drug Delivery 17 
Fig. 7. Two captured pictures from a video of a large vesicle moving in a flowing stream 
of smaller vesicles. The stills show a flexible vesicle approaching an obstacle, and rolling 
around the obstacle while adhering to it, a process encouraged by its elasticity. 
effect of nanoparticles on blood has been less well studied. Kim's31 data indicate 
that the interaction of nanoparticles with erythrocytes changes the dynamics of 
flow of both erythrocytes and particles. Chambers and Mitragotri20 found that 
nanoparticles as large as 450 nm adhered to erythrocytes, and thus remained in 
the circulation for several weeks. The percentage of latex nanospheres in the circulation 
over a period of 6 hrs was highly dependent on particle size, retention 
times decreasing with increasing diameter from 220 nm to 1100 nm. These data are 
difficult to interpret on the basis of flow, as the erythrocytes with attached nanoparticles 
are eliminated somewhat faster than the native erythrocytes. Gorodetsky and 
colleagues32 explored interactions of carboplatin (CPt) nanoparticles (formed by 
CPt interaction with fibrinogen) with the fibrin mesh caused by the induction of 
clot formation. 
18 Florence 
7. Nanoparticles with Surface Ligands 
There appear to be no rheological studies comparing surface protein decorated 
nanoparticles with the unadorned forms. Certainly, it is possible that aggregation 
may be caused by the change in surface properties and that this will in turn change 
flow patterns and perhaps masking of ligands33 as posited in Fig. 8. Nanoparticles 
are of course sensitive to the medium in which they are placed34 even in vitro when 
cell media can cause significant increases in diameter because of particle flocculation. 
We have suggested that the interaction with surface receptors of nanoparticles 
decorated with ligands is more complex than intimated in discussions of targeting 
generally.33 Figure 8 represents some of the factors: the aggregation of particles, 
the masking of ligands by this process, the detachment of ligands and the shearinduced 
removal of attached particles as discussed above. The instability of plant 
lectins, frequently used as surface proteins on nanosystems, is discussed by Gabor 
et al.35 The processes illustrated in Fig. 8 might explain some of the lack of complete 
success of targeted drug delivery. 
8. Deposition on Surfaces and Attachment to Receptors 
in Flow Conditions 
Nanoparticles in vivo flow in blood, lymph or tissue fluid at greater or lesser 
velocities, as discussed above. Deposition of particles which might occur in a 
static situation is itself a complex process, and will depend on the rugosity of the 
Aggregation and loss of 
ligand accessibility 
Repulsion Blocking by 
cleaved ligands 
n B n 
Fig. 8. Diagram illustrating variations from the ideal of a single ligand-decorated nanoparticle 
interacting with receptors spaced at an appropriate distance from the particles. The diagram 
shows the loss of ligand accessibility which would follow from the aggregation of the 
particles before interaction with the desired surface, repulsion between a particle attached 
to the receptor surface, and an approaching particle and blockage of the receptors due to 
interaction of cleaved ligands with the receptors. 
Nanoparticle Flow: Implications for Drug Delivery 19 
receiving surface.36 Particle deposition from flowing suspensions has been the subject 
of research37 which has considered not only diffusion, convection, geometrical 
interception and migration under gravity, but also the influence of tangential 
interactions. 
Patil et al.39 examined the rate of attachment of 5, 10, 15 and 20 /zm particles 
with a reconstituted P-selectin glycoprotein ligand-1 construct 19.ek.Fc. The rate of 
attachment was not affected by particle diameter. However, the shear stress required 
to set the adherent particles in motion (Sc) decreased with increasing particle diameter, 
and the rolling velocity of the 19.ek.Fc microspheres increased with increasing 
diameter. From their data, if we extrapolate the critical shear (a plot of 1 / S c is linear 
with diameter over the range 5-20 jxm), it suggests that particles below one micron 
in diameter will not be removed by shear forces. 
Usually we consider the flow of many particles in collective diffusion. The diffusion 
coefficient of a single particle and the collective diffusion coefficient coincides 
at infinite dilution, but can differ at higher concentrations.40 
Cell adhesion mediated by not one but two receptors has been considered by 
Bhatia et al.41; the analysis would also apply to decorated nanoparticles. In their 
study, the two receptors were selectin and integrin ICAM; "the state diagram" 
evolved shows the area of firm adhesion as opposed to rolling adhesion for leukocytes 
as a function of receptor densities and association rate constants. The fate of 
transport initial adhesion attachment uptake 
104- 
Distance 
(nm) 
Adhesion time 
short range interactions 
' or specific ligand e 
receptor interactions 
Fig. 9. Processes occurring in the deposition of nanoparticles in flow conditions as a function 
of the range of interaction forces (nm) and adhesion times. At the start, mass transport 
to the surface occurs, initial adhesion following through electrostatic attraction and van der 
Waals' forces. Hydrophobic interactions can play their part as well as specific receptorligand 
interactions which are short-range interactions. Drawn after Vacheethasanee and 
Marchant.38 
20 Florence 
nanoparticles in flowing blood, their adhesion, extravasation and permeation into 
tumors, thus depends on a complex of factors such as diameter, surface ligand density 
and orientation, shape, capillary diameter and rugosity, bifurcations, viscosity 
and flow gradients. 
9. Does Shape Matter? 
Nanosystems can be prepared in a variety of shapes. Nanocrystals42 are often irregular; 
there are asymmetric carbon nanotubes, and surfactant and lipid vesicles can 
be produced as discs, polyhedral structures,40,43 toroids and tubes.21,44 The vesicle 
constructs often have dimensions larger than 500 nm; it must be assumed that 
vesicles in the nanometer size range will be less affected. In these systems, shape is 
less important than membrane properties in controlling the release of encapsulated 
drug, but the flow properties of vesicular suspensions are clearly determined by 
shape and elasticity As most particulate delivery vectors have been spherical, little 
attention has been paid to the influence of shape on fate; yet it is known that the 
shape of environmental particles and fibres, for example, influences their fate and 
toxicity.45 
As discussed above, there are two different but related effects of particle flow: 
the effect of particle shape and size and characteristics on flow, as well as the effect of 
flow on flexible particles, as discussed by Bruinsma.46 With elastic vesicles, we have 
argued44 that shape matters because it affects flow and potential fate in vivo through 
extravasation for instance; elasticity also allows vesicles to be transported in vessels 
which would be blocked by solid particles. The elasticity and visco-elasticity of such 
systems may be important in differentiating them from solid nanoparticles. 
Much of the debate on whether the shape of vesicles matters, is dependent on 
the knowledge of the nature of the capillary blood supply and the forces exerted on, 
and the damage done to, vesicles as they move in capillaries.44 In studies conducted 
in our laboratories with doxorubicin loaded niosomes, 60% of the drug remained in 
the vesicles 8 hrs after intravenous administration.47 The extent to which the drug 
loss was due to diffusion or to damage is not known, but vesicles subjected to deliberate 
stress can lose considerable amounts of their payload, simply by extrusion of 
the vesicles through capillaries of reducing diameter.48 Reduction in diameter of 
systems below 1 micron will clearly reduce such stresses and allow flexible systems 
to retain their loads intact. 
Vasanthi et al.49 treated the anisotropic diffusion of oblate spheroids, explaining 
that because non-spherical molecules rotate as they translate, their motion differs 
significantly from that of a sphere. For rods, theory predicts that the diffusion 
coefficient in the direction parallel to the major axis of the rod (Dn) is twice that in 
the perpendicular direction (Di.). 
Nanoparticle Flow: Implications for Drug Delivery 21 
B IJ. • I T' ' • • • -l I 
3*M x/2 *!* 0 
Platelet angle a 
Fig. 10. The non-dimensional bond force as a function of the angle of an ellipsoidal platelet 
passing through zero when the platelet is 90° to the surface. From Mody et a/.50 
There are few studies which have considered the motion of ellipsoidal particles 
near a plane wall, although this is relevant to platelet flow and adhesion to the walls 
of vessels. Mody and colleagues50 have addressed the issue, observing the effects 
of shear stress on platelet adhesion. Platelets, unlike leukocytes, do not roll but 
display a flipping motion in the direction of flow, due to their flattened ellipsoidal 
structure. The bond force between the ellipse and the surface is dependent on the 
platelet angle as defined in Fig. 10. 
Flexible systems such as vesicles have been widely studied, while being forced 
under pressure in capillaries smaller than the vesicle diameter. The elasticity of 
the membranes can be estimated from the extent of deformation. Vesicle flow in 
linearly forced motion has been followed. Flexible vesicles adjust their shape to 
equilibrate the applied force51; locally in some cases, two-dimensional flow of lipids 
in the vesicle membrane occurs,52 clearly influencing the position of the embedded 
surface ligands. 
There are many nanoparticulates which are produced in non-spherical forms, 
hence the transport properties of asymmetric particles is important.53 
10. Speculations on Flow and the EPR Effect 
Erythrocyte velocity in normal vessels depends on vessel diameter (see Fig. 4 
above), but there is no such dependence in tumors (Fig. 11), even though flow 
may be an order of magnitude slower. According to Jain,18,52 "to reach cancer cells 
in a tumor, a blood-borne therapeutic molecule, particle or cell must make its way 
22 Florence 
0.8 
0.7 
•t 0.5 
0.4 
> 
i 0.2 
o.i H 
0.0 
MCalV J U»7 
i r*—I—1 r——i 1 I i ""—> r 1 1 T" 
0 10 20 30 40 SO 60 70 0 10 20 30 40 50 60 70 
Tumor Vessel Diameter Qua) 
Fig. 11. Diagram from Jain18 showing the lack of a clear relationship between erythrocyte 
velocity and tumor vessel diameter in two tumor types, MCalV and U87. The low and 
variable velocities compared to those shown in Fig. 4 are evident. 
into blood vessels of the tumor and cross the vessel wall into the interstitium and 
finally migrate through the interstitium". While blood flow is reduced in tumor 
vessels, nonetheless cancer cells have been reported to compress tumor vessels and 
this will have consequences on fluid flow.54 This is highly relevant to the enhanced 
permeation and retention effect (EPR) which allows entry of macromolecules into 
tumors from spaces in the ill-formed tumor vasculature.55 Access of nanoparticles 
to tumors is equally important and must be critically size-dependent. 
In convective flow, stable colloidal particles may be captured by the process of 
hydrodynamic bridging,52,56 events which may be relevant to the first process in 
the enhanced permeation and retention (EPR) effect. At high velocities but in the 
low Re regime, hydrodynamic forces acting on the particles at an entrance to a pore 
(or a defect in a tumor vessel) may overcome colloidal repulsive forces and result 
in flocculation of the particles and the plugging of the pore. The effects of velocity, 
particle concentration, and the ratio of pore size to particle size (the aspect ratio) on 
retention by hydrodynamic bridging have been studied. The effect of velocity on 
retention by bridging is opposite to that of retention by deposition. There is a critical 
flow velocity necessary for particle bridging to occur, a function of the net colloidal 
interparticle and particle—porous medium repulsion that must be overcome by the 
hydrodynamic forces for bridging to occur. Figure 12 demonstrates the effect for an 
aspect ratio of 3.7 (220 nm particles) 
11. Intra-tumoral Injection 
Direct injection of delivery systems into tumors has both been a mode of experimental 
and clinical drug delivery. Solutions allow the drugs to diffuse or leach out 
Nanoparticle Flow: Implications for Drug Delivery 23 
f l 
v^ o- 
• f 
• * 
Fig. 12. Particle behavior prior to entry to a pore of radius, rp: (a) a discrete nanoparticle, 
(b) aggregate, (c) individual particles converging on the pore opening demonstrating 
hydrodynamic bridging, as discussed by Ramachandran.56 We speculate that events such 
as bridging might occur during entry of nanoparticles into tumors through fenestrations in 
the tumor capillary blood supply, aspects of the enhanced permeation and retention effect. 
of the tumor, especially through the needle track, whereas suspensions might allow 
some greater residence time. Viral vectors have been administered by intra-tumoral 
injection.57 To decrease the extent of viral dissemination into the systemic circulation, 
a viscous alginate solution was used as the viral vehicle. However, transgene 
expression was not increased perhaps because, as the authors speculate, the diffusion 
of the virus is reduced by the viscous medium once in situ. The transport 
of particles of viral dimensions requires, according to Higuchi et al.,16 convective 
rather than diffusional transport. "The early transport of colloids into the vascular 
and lymphatic vessels relies largely on an extracellular pathway which depends 
on convective transport (i.e. solvent drag)". "Thus the particle uptake in the period 
immediately after injection is relatively insensitive to particle size; it is expected 
that viruses will be carried in the tissue towards lymphatics and microvessels with 
great efficacy leading to enhanced escape compared with the relatively low levels" 
for 1 and 0.4 /xm particles.16 The question of how resistance to convective transport 
in the interstitial space (the interstitial fluid plus the extracellular matrix) has been 
considered at least for molecules.58 Clearly, the spacing between the cells or between 
fibres will be a significant factor in determining the size cut-off for transport. 
12. Conclusions 
This phenomenological survey of possible factors affecting the flow and hence the 
mass transport of nanoparticles has explored a range of scenarios. It is by no means 
a comprehensive survey, but there is sufficient in the literature to stimulate further 
analyses to provide a better overall prediction of the influence of particle characteristics, 
particularly, diameter and surface nature, shape and flexibility on delivery 
and targeting to remote sites in the body. Conf ocal microscopy and other techniques 
24 Florence 
will allow experimental study of nanoparticles so that their movement and fate can 
be studied in a variety of tissues. Atomic force microscopy allows measurement of 
forces of interaction of particles with cells and receptors to aid a more quantitative 
approach. However, it is wrong to underestimate the challenges ahead if nanoparticulate 
carriers are to be designed to overcome the various biological barriers and 
survive transit in the conduits of capillary blood or lymph, extravasation and tissue, 
and subsequently intracellular transport.59 One cannot help but conclude that as 
many properties including flow are dictated by particle diameter, one of the most 
important strategies is to ensure the maintenance of particle stability in vivo. 
References 
1. El-Sayed M, Kiani MF, Naimark MD, Hikal AH and Ghandehari H (2001) Extravasation 
of poly(amidoannine) (PAMAM) dendrimers across microvascular network endothelium. 
Pharm Res 18:23-28. 
2. Health and Safety Executive, Health Effects of particles produced for nanotechnologies. 
Sudbury, UK, pp. 1-37. 
3. Hussain N, Jaitley V and Florence AT (2001) Recent advances in the understanding of 
uptake of microparticulates across the gastrointestinal lymphatics. Adv Drug Del Rev 
50:107-142. 
4. Florence AT (1997) The oral absorption of micro- and nanoparticulates: Neither exceptional 
nor unusual. Pharm Res 14:259-266. 
5. Florence AT and Hussain N (2001) Transcytosis of nanoparticle and dendrimer delivery 
systems: Evolving vistas. Adv Drug Del Rev 50 (Suppl 1):S69-S89. 
6. Fortina P, Kricka LJ, Surrey S and Grodzinski P (2005) Nanobiotechnology: The 
promise and reality of new approaches to molecular recognition. Trends Biotechnol 23: 
168-173. 
7. Warheit DB, Laurence BR, Reed KL, Roach DH, Reynolds GA and Webb TR (2004) 
Comparative pulmonary toxicity assessment of single-wall carbon nanotubes in rats. 
Toxicol Sci 77:117-125. 
8. Sugihara-Seki M and Skalak R (1997) Asymmetric flows of spherical particles in a 
cylindrical tube. Biorheology 34:155-159. 
9. Wang H and Skalak R (1969) Viscous flow in a cylindrical tube containing a line of 
spherical particles. / Fluid Mech 38:75-96. 
10. Jani P, Halbert GW, Langridge J and Florence AT (1989) The uptake and translocation of 
latex nanospheres and microspheres after oral-administration to rats.}Pharm Pharmacol 
41:809-812. 
11. Nasseri B and Florence AT (2003) Microtubules formed by capillary extrusion and 
fusion of surfactant vesicles. Int} Pharm 266:91-98. 
12. Rowland RES, Taylor PW and Florence AT (2005) / Drug Del Sci Tech 
13. Ruenraroengsak P, Hartell N and Florence AT (2005) unpublished. 
14. Fokin AA, Robicsek F and Masters TN (2000) Transport of viral-size particulate matter 
after intravenous versus intralymphatic entry. Microcirculation 7:357-365. 
Nanoparticle Flow: Implications for Drug Delivery 25 
15. Fokin AA, Robicsek F, Masters TN, Schmid-Schonbein GW and Jenkins SH (2000) 
Propagation of viral-size particles in lymph and blood after subcutaneous inoculation. 
Microcirculation 7:193-200. 
16. Higuchi M, Fokin A, Masters TN, Robicsek F and Schmid-Schonbein GW (1999) Transport 
of colloidal particles in lymphatics and vasculature after subcutaneous injection. 
JAppl Physiol 86:1381-1387. 
17. Ilium L, Davis SS, Wilson CG, Thomas NW, Frier M and Hardy JG (1982) Blood clearance 
and organ deposition of intravenously administered colloidal particles. The effects of 
particle size, nature and shape. Int ] Pharm 12:135-146. 
18. Jain RK (2001) Delivery of molecular medicine to solid tumors: Lessons from in vivo 
imaging of gene expression and function. / Control Rel 74:7-25. 
19. Silebi CA and DosRamos JG (1989) Separation of submicrometer particles by capillary 
hydrodynamic fractionation (CHDF). / Coll Interf Sci 130:14-24. 
20. Chambers E and Mitragotri S (2004) Prolonged circulation of large polymeric nanoparticles 
by non-covalent adsorption on erythrocytes. / Control Rel 100:111-119. 
21. Nunez ADR, Pinto R and Paredes VME (2002) Viscosity minimum in bimodal concentrated 
suspensions under shear. Eur Phys } E 9:327-334. 
22. Ding Y and Wen D (2005) Particle migration in a flow of nanoparticle suspensions. 
Powder Technol 149:84-92. 
23. Odde D (1998) Diffusion inside microtubules. Eur Biophys J 27:514-520. 
24. Sinton D (2004) Microscale flow visualization. Microfluid Nanofluid 1:2-21. 
25. Chiu J-J, Chen C-N, Lee P-L, Yang CT, Chuang HS and Chien SUS (2003) Analysis 
of the effect of disturbed flow in monocytic adhesion to endothelial cells. / Biomech 
26:1883-1895. 
26. Shankar A, Loizidou M, Burnstock G and Taylor I (1999) Noradrenaline improves the 
tumour to normal blood flow ratio and drug delivery in a model of liver metastases. 
Br } Surgery 86:453^57. 
27. Goldberg JA, Murray T, Kerr DJ, Willmott N, Bessent RG, McKillop JH and McCardle 
CS (1991) The use of angiotensin II as a potential method of targeting cytotoxic microspheres 
in patients with intrahepatic tumours. Br J Cancer 63:308-310. 
28. Zhang Z, Kleinstreuer C, Donohue JF and Kim CS (2005) Comparison of micro- and 
nano-size particle depositions in a human upper airway model. Aerosol Sci 36:211-233. 
29. Shi HKC, Zhang Z and Kim CS (2004) Nanoparticle transport and deposition in bifurcating 
tubes with different inlet conditions. Phys Fluids 16:2199-2213. 
30. James SC and Chrysikopoulos CV (2004) Dense colloid transport in a bifurcating fracture. 
/ Coll Interf Sci 270:250-254. 
31. Kim D, El-Shall H, Dennis D and Morey T (2005) Interaction of PLGA nanoparticles 
with human blood constituents. Coll SurfB 40:83-91. 
32. Gorodetsky R, Peylan-Ramu N, Reshef A, Gaberman E, Levdansky L and Marx G 
(2005) Interactions of carboplatin with fibrin(ogen), implications for local slow release 
chemotherapy. / Control Rel 102:235-245. 
33. Florence AT (2005) Issues in oral nanoparticle drug carrier uptake and targeting. / Drug 
Targ 12:65-70. 
26 Florence 
34. Singh B, Hussain N, Sakthivel T and Florence AT (2003) Effect of physiological media on 
the stability of surface-adsorbed DNA-dendron-gold nanoparticles. / Pharm Pharmacol 
55:1635-1640. 
35. Gabor F, Bogner E, Weissenboeck A and Wirth M (2004) The lectin-cell interaction and its 
implications to intestinal lectin-mediated drug delivery. Adv Drug Del Rev 56:459-480. 
36. Adamczyk Z, Siwek B, Jaszczolt K and Weronski P (2004) Deposition of latex particles 
at heterogeneous surfaces. Colloids Surface A: Physicochem Eng Aspects 249:95-98. 
37. Adamczyk Z (1989) Particle transfer and deposition from flowing colloid suspensions. 
Coll Surf 35:283-308. 
38. Vacheethasanee K and Marchant RE (2000) Non-specific staphylococcus epidermidis 
adhesion: Contribtuions of biomaterial hydrophobicity and charge, in An, YH, 
Friedman RJ (eds.) Handbook of Bacterial Adhesion: Principles, Methods and Applications. 
Humana Press, Totowa, NJ, pp. 73-90. 
39. Patil VRS, Campbell CJ, Yun YH, Slack SM and Goettz DJ (2001) Particle diameter 
influences adhesion under flow. Biophys J 80:1733-1743. 
40. Bowen WR and Mongruel A (1998) Calculation of the collective diffusion coefficient of 
electrostatically stabilised colloidal particles. Coll Surface A 138:161-172. 
41. Bhatia SK, King MR and Hammer DA (2003) The state diagram for cell adhesion mediated 
by two receptors. Biophys J 84:2671-2690. 
42. Akerman ME, Chan WC, Laakkonen P, Bhatia SN and Ruoslahti E (2002) Nanocrystal 
targeting in vivo. Proc Natl Acad Sci USA 99:12617-12621. 
43. Uchegbu IF, Schatzlein A, Vanlerberghe GMN and Florence AT (1997) Polyhedral nonionic 
surfactant vesicles. J Pharm Pharmacol 49:606-610. 
44. Florence AT, Nasseri B and Arunothyanun P (2004) Does shape matter? Spherical, 
polyhedral and tubular vesicles, in Sonke S (ed.) Carrier-based Drug Delivery. American 
Chemical Society, Washington, pp. 75-84. 
45. Schins RP (2002) Mechanisms of genotoxicity of particles and fibers. Inhal Toxicol 14: 
57-78. 
46. Bruinsma R (2005) Rheology and shape transitions of vesicles under capillary flow. 
Physica A 234:249-270. 
47. Uchegbu IF, Double JA, Turton JA and Florence AT (1995) Distibution, metabolism and 
tumoricidal activity of doxorubicin administered in sorbitan monostearate (Span 60) 
niosomes in the mouse. Pharm Res 12:1019-1024. 
48. Nasseri B and Florence AT (2003) Some properties of extruded non-ionic surfactant 
micro-tubes. Int f Pharm 254:11-16. 
49. Vasanthi R and Bhattacharyya S (2005) Anisotropic diffusion of spheroids in liquids: 
Slow orientational relaxation of the oblates. / Chem Phys 116:1092-1096. 
50. Mody NA, Lomakin O, Doggett TADTG and King MR (2005) Mechanics of transient 
platelet adhesion to von Willebrand factor under flow. Biophys J 88:1432-1443. 
51. Kern N and Fourcade B (1999) Vesicles in linearly forced motion. Europhys Lett 
46:262-267. 
52. Nasseri B and Florence AT (2005) The relative flow of the walls of phospholipid tethers. 
Int J Pharm 298:372-377. 
Nanoparticle Flow: Implications for Drug Delivery 27 
53. Naess SN and Elgsaeter A (2005) Transport properties of non-spherical nanoparticles 
studied by Brownian dynamics: Theory and numerical simulations. Energy 30:831-844. 
54. Padera TP, Stoll BR, Tooredman JB, Capen D, di Tomaso E and Jain RK (2004) Cancer 
cells compress intratumour vessels. Nature 427:695. 
55. Maeda H (2001) The enhanced permeability and retention (EPR) effect in tumor vasculature: 
the key role of tumor-selective macromolecular drug targeting. Adv Enzyme 
Regul 41:189-207. 
56. Ramachandran VV, Venkatesan R, Tryggvason G and Scott FH (2000) Low Reynolds 
Number Interactions between Colloidal Particles near the Entrance to a Cylindrical 
Pore. / Coll Interf Sci 229:311-322. 
57. Wang Y, Hu JK, Krol A, Li YP, Li CY and Yuan F (2003) Systemic dissemination of viral 
vectors during intratumoral injection. Mol Cancer Ther 2:1233-1242. 
58. McGuire S and Yuan F (2001) Quantitative analysis of intratumoral infusion of color 
molecules. Am J Physiol Heart Circ Physiol 281:H715-H721. 
59. Jones AT, Gumbleton M and Duncan R (2003) Understanding endocytic pathways and 
intracellular trafficking; a prerequisite for effective design of advanced drug delivery 
systems. Adv Drug Del Rev 55:1353-1357. 
This page is intentionally left blank
3 
Polymeric Nanoparticles as Drug 
Carriers and Controlled Release 
Implant Devices 
SM Moghimi, E Vega, ML Garcia, 
OAR Al-Hanbali and KJ Rutt 
1. Introduction 
Polymeric nanoparticles are submicron size entities, often ranging from 10-1000 nm 
in diameter, and are assembled from a wide variety of biodegradable (e.g. albumin, 
chitosan, alginate) and non-biodegradable polymers (Tables 1 and 2). The most 
active area of research using polymeric nanoparticles is in controlled delivery of 
pharmaceuticals following parenteral, oral, pulmonary, nasal, and topical routes 
of administration.1-6 Indeed, therapeutic agents can be encapsulated, covalently 
attached, or adsorbed onto such nanocarriers. These approaches can easily overcome 
drug solubility issues; this is particularly important as a significant proportion 
of new drug candidates arising from high-throughput screening initiatives are 
water insoluble. Polymeric nanoparticles, however, differ from nanosuspensions 
of drugs which are sub-micron colloidal dispersions of pure particles of drug that 
are stabilized by surfactants.7 By virtue of their small size and by functionalizing 
their surface with polymers and appropriate ligands, polymeric nanoparticles can 
also be targeted to specific cells and locations in the body.1,3'5'8-10 Thus, polymeric 
nanoparticles may overcome stability issues for certain drugs and minimize druginduced 
side effects. The extent of drug encapsulation/incorporation, as well as 
29 
30 Moghimi etal. 
the release profile from polymeric nanocarriers, however, depends on the polymer 
type and its physicochemical properties, the particle size and its morphology (e.g. 
solid nanospheres as opposed to polymeric nanocapsules).4 In addition, depending 
on the polymer characteristics, polymeric nanocarriers can also be engineered 
in such a way that they can be activated by changes in the environmental pH, 
chemical stimuli, or temperature.1112 Such modifications offer control over particle 
integrity, drug delivery rates, and the location of drug release, for example, 
within specific organelles. For instance, nanoparticles made from poly(lactide-coglycolide), 
PLGA, can escape the endo-lysosomal compartment within minutes 
of internalization in intact cells and reach the cytosol.12 This is due to the selective 
reversal of the surface charge of nanoparticles from the anionic to the cationic state in 
endo-lysosomes, resulting in a local particle-membrane interaction with subsequent 
cytoplasmic release. This is an excellent approach for channelling antigens into the 
highly polymorphic MHC class-I molecules of macrophages and dendritic cells 
for subsequent presentation to CD8+ T lymphocytes. Other applications include 
cytoplasmic release of plasmid vectors and therapeutic agents (e.g. for combating 
cytoplasmic infections and for slow cytoplasmic release of drugs that act on nuclear 
receptors). 
Polymeric nanoparticles are also beginning to make a significant impact on 
global pharmaceutical planning (life-cycle management) and market intelligence. 
For example, due to imminent expiration of patents, pharmaceutical companies 
may launch follow-up or nano-formulated versions of a product to minimize 
generic threats to best-selling medicines. This could lead to an extension of as much 
as 20 years from a new patent on the nanoparticulate formulation of the drug. 
By coalescing certain polymeric nanoparticles carefully from an aqueous 
suspension, shape retentive hydrogels can be formed to erode partially or 
completely.1113 Drugs and macromolecules may be trapped within interstitial 
spaces between particles during aggregate formation. Thus, hydrogel nanoparticles 
have potential as controlled release implant devices following local administration 
or implantation, and may also serve as tissue engineering scaffolds with concurrent 
morphogenic protein release. 
This article will briefly review some of the most commonly used laboratory 
scale methods for the production of polymeric nanoparticles and drug encapsulation 
procedures. The importance of the nanometre scale size range and surface 
engineering strategies for site-specific targeting of polymeric nanoparticles, following 
different routes of administration, are also discussed. 
2. Nanoparticle Engineering 
Polymeric nanoparticles are usually prepared either directly from preformed 
polymers such as aliphatic polyesters (Table 1) and block copolymers (Table 2), 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 31 
Table 1 Chemical properties of some commonly used aliphatic polyesters in nanoparticle 
engineering. 
Polymer Type Melting Point (°C) Glass Transition Resorption Time 
Temperature (°C) (Months) 
DL-PLA Amorphous 50-60 12-16 
PGA 220-230 35^0 6-12 
DL-PLGA (50/50) Amorphous 45-50 1-2 
DL-PLGA (75/25) Amorphous 45-50 4-5 
PCL 55-65 (-65)-(-60) >24 
DL-PLA: poly(L-lactide); PGA: poly(glycolide); DL-PLGA: poly(DL-lactide-co-glycolide); PCL: poly- 
.-caprolactone. 
Table 2 Selected examples of block copolymers for production of biodegradable 
nanospheres. 
PLA-poly(ethyleneglycol),PLA-PEG 
MonomethoxyPEG-poly(alkylcyanoacrylate) 
Poly(poly(ethyleneglycol)cyanoacrylate-co-hexadecylcyanoacrylate) 
Poly(ethyleneoxide-b-sebacicacid) 
Poly(phosphazene)-poly(ethyleneoxide) 
poly(2-methyloxazoline)-b-poly(dimethylsiloxane)-b-poly(2-methyloxazoline) 
or by polymerization of monomers.4 Commonly used methodologies include 
the solvent evaporation,14-15 the spontaneous emulsification/ solvent diffusion,16 
nanoprecipitation or solvent displacement17'18 and emulsion polymerization 
techniques.19-21 The method of choice depends on the polymer and the drug 
type, as well as the required particle size distribution and polydispersity 
indices. However, some polymers, such as comb-like polyesters, the di-block 
copolymer poly(ethylene oxide-b-sebacic acid) and tri-block copolymer poly(2- 
methyloxazoline)-fr-poly(dimethylsiloxane)-fr-poly(2-methyloxazoline) can spontaneously 
form stable nanoparticles (core-shell type nanospheres).22-24 
In the solvent evaporation method, the polymer is simply dissolved together 
with the drug in an organic solvent and the mixture is then emulsified to form either 
an oil-in-water nanoemulsion (for encapsulation of hydrophobic drugs) or waterin-
oil nanoemulsion (for encapsulation of hydrophilic drugs) using suitable surfactants. 
Nanoparticles are then obtained following evaporation of the solvent and 
can be concentrated by filtration, centrifugation or lyophilization. The spontaneous 
emulsification/solvent diffusion method is a modified version of the solvent evaporation 
technique, which utilizes a water-soluble solvent (e.g. methanol or acetone) 
along with a water-insoluble one such as chloroform. As a result of the spontaneous 
32 Moghimi etal. 
diffusion of the water-soluble solvent into the water-insoluble phase, an interfacial 
turbulence is created leading to the formation of nanoparticles. Nanoprecipitation, 
however, is a versatile and simple method. This is based on spontaneous formation 
of nanoparticles during phase separation (the Marangoni effect), which is induced 
by slow addition of the diffusing phase (polymer-drug solution) to the dispersing 
phase (a non-solvent of the polymers, which is miscible with the solvent that solubilizes 
the polymer). The dispersing phase may contain surfactants. Depending 
on the solvent choice and solvent/non-solvent volume ratio, this method is suitable 
for encapsulation of both water-soluble and hydrophobic drugs, as well as 
protein-based pharmaceuticals.17'18 
In emulsion polymerization, the monomer is dispersed into an aqueous phase 
using an emulsifying agent. The initiator radicals are generated in the aqueous 
phase and they diffuse into the monomer-swollen micelles. Anionic polymerization 
in the micelles is then initiated by the hydroxyl ions of water. Chain transfer 
agents are abundant and termination occurs by radical combination. The size and 
molecular masses of nanoparticles are dependent on the initial pH of the polymerization 
medium.20 Drugs are incorporated during the polymerization step or can 
be adsorbed into the nanosphere surface afterwards. The addition of cyclodextrins 
to the polymerization medium can promote the encapsulation of poorly watersoluble 
drugs.25 Depending on the monomer used, some drugs can also initiate the 
polymerization step, resulting in the covalent attachment of drug molecules to the 
nanospheres. For instance, photosensitizers such as naphthalocyanines, can initiate 
the polymerization of alkylcyanoacrylates.26 
A number of specialized approaches (e.g. dialysis, salting-out, supercritical 
fluid technology, denaturation, ionic interaction, ionic gelation, and interfacial 
polymerization) have also been described for the preparation of polymeric 
nanoparticles, based on the choice of the starting material and the biological 
needs.4'27-32 
2.1. Drug release mechanisms 
The release profile of drugs from nanoparticles depends on the physicochemical 
nature of the drug molecules as well as the matrix.4'16'28,33-36 Factors include mode of 
drug attachment and/or encapsulation (e.g. surface adsorption, dispersion homogeneity 
of drug molecules in the polymer matrix, covalent conjugation), the physical 
state of the drug within the matrix (such as crystal form), and parameters controlling 
matrix hydration and/or degradation. Generally, rapid release occurs by desorption, 
where the drug is weakly bound to the nanosphere surface. If the drug is 
uniformly distributed in the polymer matrix, the release occurs either by diffusion 
(if the encapsulated drug is in crystalline form, the drug is first dissolved locally 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 33 
and then diffuses out) or erosion of the matrix, or a combination of both mechanisms. 
Erosion can be further subdivided into either homogeneous (with uniform 
degradation rates throughout the matrix) or heterogeneous (where degradation is 
confined at the surface) processes. Parameters such as polymer molecular weight 
distribution, crystallinity, hydrophobicity/hydrophilicity, melting and glass transition 
temperature, polymer blends and prior polymer treatment (e.g. oxygen-plasma 
treatment) all control the extent of matrix hydration and degradation. For instance, 
in the case of aliphatic polyesters, their degradation time is shorter for low molecular 
weight polymers, more hydrophilic polymers, more amorphous polymers and 
copolymers with high glycolide content (Table 1). 
3. Site-specific Targeting with Nanoparticles: Importance 
of Size and Surface Properties 
Numerous articles have recently discussed the importance of nanoparticle size and 
surface characteristics in controlling their biodistribution, following different routes 
of administration.1 ~3/5 Only a brief overview is provided here. 
Following intravenous injection, liver (Kupffer cells) and spleen (marginal zone 
and red pulp) macrophages clear polymeric nanoparticles rapidly from the blood 
circulation.1 Opsonization, which is surface deposition of blood opsonic factors 
such as fibronectin, immunoglobulins, C-reactive and certain complement proteins, 
often aid particle recognition by these macrophages. Indeed, the propensity 
of macrophages of the reticuloendothelial system for rapid recognition and 
clearance of particulate matter has provided a rational approach to macrophagespecific 
targeting with nanoparticles (e.g. for the treatment of obligate intracellular 
microorganisms, delivery of toxins for macrophage killing, and diagnostic agents).1 
However, the rapid sequestration of nanoparticles by macrophages in contact with 
blood is problematic for the efficient targeting of polymeric nanoparticles to nonmacrophage 
sites. Thus, inherent in nanoparticle design is the precision surface 
manipulation and engineering with synthetic polymers; this affords control over 
nanoparticle interaction and fate within biological systems. There are numerous 
examples where the surface of nanocarriers is carefully assembled with projected 
"macromolecular hairs" made from poly(ethyleneglycol), PEG, or its derivatives 
(e.g. methoxyPEG-albumin, PLA-PEG) or other related polymers [e.g. block 
copolymers such as selected poloxamers and poloxamines, poly(phosphazene)- 
poly(ethyleneoxide)].3,5 This is achieved either during the particle assembly procedures 
or polymerization step, or post particle manufacturing. This strategy 
suppresses macrophage recognition by an array of complex mechanisms, which 
collectively achieve reduced protein adsorption and surface opsonization. Therefore, 
such entities, provided that they are below 150 nm in size, exhibit prolonged 
34 Moghimi et al. 
residency time in the circulation, and are referred to as "stealth" or "macrophageevading" 
nanoparticles.1,5 The efficiency of the "macrophage-evading" process is 
dependent on polymer type and its surface stability, reactivity, and physics (e.g. 
surface density and assumed conformation).5 Prolonged circulation properties are 
ideal for slow or controlled release of therapeutic agents in the blood to treat 
vascular disorders. Long circulating polymeric nanoparticles may have application 
in vascular imaging too (e.g. detection of vascular bleeding or abnormalities). 
Long-circulating nanoparticles can also escape from vasculature and this is normally 
restricted to sites where the capillaries have open fenestration or when the 
integrity of the endothelial barrier is perturbed by inflammatory processes or by 
tumor growth.5 However, extravasated nanoparticles, as in tumour interstitium, 
distribute heterogeneously in perivascular clusters that do not move significantly; 
these particles may therefore act as depot systems, particularly for the sustained 
release of antiangiogenic agents, and to some extent, for drug delivery to multidrug 
resistant tumors (e.g. by co-encapsulation of both anticancer drugs and the competitive 
inhibitors of active drug efflux pumps).1 The surface of long-circulating 
nanoparticles is also amenable for modification with targeting ligands. Such entities 
can navigate capillaries and escape routes in search of signature molecules 
expressed by the target; this process is often referred to as "active targeting".1-5 
For example, certain cancer cells express folate receptors and these receptors have 
the ability to endocytose stealth nanoparticles that are decorated with folic acid. 
Delivery of anti-cancer agents to tumor cells by such means could overcome the 
possibility of multi-drug resistance.1,37 
Non-deformable "stealth" nanoparticles, however, are prone to splenic filtration 
at interendothelial cell slits, if their size exceeds that of the width of the cell 
slits (200-250 nm).38,39 Indeed, these "splenotropic" vehicles can deliver their cargo 
efficiently to the red-pulp regions of the sinusoidal spleen. Activated or stimulated 
macrophages are also known to rapidly phagocytose stealth nanoparticles; 
stealth nanospheres may therefore have applications as diagnostic/imaging tools 
for the identification of stimulated or newly recruited hepatic macrophages.40 Such 
diagnostic procedures may prove useful for patient selection or for monitoring 
the progress of treatment with long-circulating nanoparticles carrying anti-cancer 
agents, thus minimizing damage to hepatic macrophages.41 
Polymeric nanospheres can also target endothelial cells on the bloodbrain 
barrier. For instance, following intravenous injection polysorbate 80-coated 
poly(alkylcyanoacrylate), PACA, nanospheres attract apolipoprotein E from the 
blood, thus mimicking low density lipoprotein (LDL) and become recognizable 
by LDL receptors expressed by the blood-brain barrier endothelial cells.10 Another 
related example is PEG-coated PACA nanoparticles, with the ability to localize 
mainly in the ependymal cells of the choroid plexus and the epithelial cells of pia 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 35 
region and the ventricles of the mouse and the rat brain.42 The molecular basis of 
this deposition pattern remains to be unravelled. 
Others have administered nanoparticles directly to pathological sites for 
optimal biological performance.43 One example is intramurally delivered PLGA 
nanoparticles to an injured artery following angioplasty, using a cardiac infusion 
catheter. Here, nanoparticles penetrate the dilated arterial wall under pressure and 
once the pressure is released, the artery returns to its normal state resulting in particle 
immobilization in the arterial wall, where they may act as a sustained release 
system for drugs and genetic materials.43 Again, particle size is an important parameter; 
the smaller the size, the greater the arterial deposition and cellular entry, as 
well as lower inflammatory responses. 
Polymeric nanospheres also provide intriguing opportunities for lymphatic 
drug delivery, as well as for diagnostic imaging of the lymphatic vessels and their 
associated lymph nodes when injected interstitially.44 The extent of lymphatic delivery 
and lymph node localization of nanospheres depends on their size and surface 
characteristics. For instance, hydrophilic nanoparticles, in the size range of 
30-100 nm, as opposed to their hydrophobic counterparts, repulse each other and 
interact poorly with the ground substance of the interstitium and drain rapidly into 
the initial lymphatics through patent junctions in the lymphatic capillaries.45,46 The 
drained particles are conveyed to the nodes via the afferent lymph. Macrophages 
of medullary sinuses and paracortex are mainly responsible for particle capture 
from the lymph, but this also depends on nanoparticle surface properties. Larger 
nanospheres (>150nm), however, are retained at interstitial sites for prolonged 
periods of time and may therefore act as sustained release systems for drugs and 
antigens.47,48 For example, large-sized PLGA particles can provide antigen release 
over weeks and months following continuous or pulsatile kinetics. By mixing particle 
types with different degradation and pulsatile release kinetics, multiple discrete 
booster doses of encapsulated antigens can be provided after a single administration 
of the formulation (e.g. 1-2 and 6-12 months).48 An alternative approach is the use 
of nanoparticle hydrogels for slow and local antigen release. For example, by controlling 
the ionic strength of the dispersion medium, monodisperse nanoparticles of 
poly-2-hydroxyethylmethacrylate, poly(HEMA), and poly[HEMA-co-methacrylic 
acid] coalesce together to form a shape retentive hydrogel suitable for interstitial 
implantation.13 Macromolecules may be trapped between the particle aggregates 
and their release is controlled by a combination of diffusion (larger particles packed 
together have larger spaces in the lattice, and this allows for faster diffusion) and 
erosion (arising from aggregates that contain particles with methacrylic acid).13 
Nanoparticles that erode from the aggregate are drained into the lymphatic system 
and may be retained by the regional nodes. Similarly, by controlling the inherent 
physical attractive forces between model polystyrene nanoparticles, ordered lattices 
36 Moghimi et al. 
Fig. 1. Scanning electron micrographs of uncoated and surface-modified polystyrene 
nanoparticles. Due to surface hydrophobicity uncoated nanospheres (A), 350 nm in size, 
tend to aggregate. By controlling the physical attractive forces between the nanoparticles (by 
surface coating with an appropriate concentration of a block copolymer), ordered structures 
are formed and these can be deposited onto the surface of large microspheres (B). 
can be deposited on the surface of very large microspheres (Fig. 1). Following subcutaneous 
localization, surface adsorbed nanospheres may gradually detach from 
the parent microsphere and gain entry into the lumen of the lymphatic capillaries. 
Polymeric nanoparticles also have numerous applications following oral delivery. 
Evidence suggests that the adsorption of particulates in the intestine following 
oral administration take place at the Peyer's patches.49-50 The epithelial cell 
layer overlying the Peyer's patches contains specialized M cells. These cells can 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 37 
sample particles from the lumen and transport them to the underlying macrophages 
and dendritic cells. Indeed, numerous studies have confirmed protective immunity 
induced by mucosal immunization with PACA, PLGA and chitosan based particulate 
systems.3,32,48'50-53 Part of the success is due to the encapsulation of antigens in 
polymeric particulate systems, which provides better protection for the antigen during 
intestinal transit. The immune outcomes have included mucosal (secretory IgA) 
and serum antibody (IgG and IgM) responses, as well as systemic cytotoxic T lymphocyte 
responses in splenocytes. Induction of an appropriate immune response 
following oral administration depends primarily on factors that affect uptake and 
particle translocation by M cells. These include particle size, dose, composition, and 
surface chemistry, as well as the region of the intestine where particles are taken up, 
membrane recycling from intracellular sources and the species.50 Tolerance to orally 
administered microparticulate encapsulated antigens is another potential outcome, 
but it has received little attention. 
The bioavailability of some drugs can be improved after oral administration 
by means of polymeric nanoparticles.54-57 This is a reflection of drug protection 
by the nanoparticle against hostile conditions of the gastrointestinal tract, as well 
as the mode of nanoparticle interaction with mucosal layers. However, the bioadhesive 
properties of nanoparticles may vary with their size and surface characteristics 
(e.g. surface charge, surface polymer density and conformation), as well as 
the location and type of the mucosal surface in the gastrointestinal tract. Similarly, 
improved drug bioavailability has also been reported following ocular administration 
with PLA, PACA, poly(butylcyanoacrylate) and Eudragit nanoparticles.6,58-61 
For example, loading of tamoxifen in PEGylated nanoparticles proved successful 
in the treatment of autoimmune uveortinitis following intraocular injection.59 
Interaction of surface-modified polymeric nanoparticles with nasal associated lymphoid 
tissue and their transport across nasal mucosa have also received attention, 
particularly with respect to peptide-based pharmaceuticals and antigen 
delivery.53,62 
4. Conclusions 
Polymeric nanoparticles are promising vehicles for site-specific and controlled 
delivery of therapeutic agents, following different routes of administration and 
these trends seem to continue with advances in materials and polymer chemistry 
and pharmaceutical nanotechnology. However, nanoparticles do not behave similarly; 
their encapsulation capacity, drug release profile, biodistribution and stability 
vary with their chemical makeup, morphology and size. Inherently, nanosphere 
design and targeting strategies may vary according to physiological and therapeutic 
needs, as well as in relation to the type, developmental stage and location of 
38 Moghimietal. 
the disease. Attention should also be paid to toxicity issues that may arise from 
nanoparticle administration and the release of their polymeric contents and degradation 
products. These issues are discussed elsewhere.1,63~66 
References 
1. Moghimi SM, Hunter AC and Murray JC (2005) Nanomedicine: Current status and 
future prospects. FASEB ] 19:311-330. 
2. Panyam J and Labhasetwar V (2003) Biodegradable nanoparticles for drug and gene 
delivery to cells and tissue. Adv Drug Del Rev 55:329-347. 
3. Vauthier C, Dubernet C, Fattal E, Pinto-Alphandary H and Couvreur P (2003) Poly 
(alkyleyanoacrylates) as biodegradable materials for biomedical applications. Adv Drug 
Del Rev 55:519-548. 
4. Soppimath KB, Aminabhavi TM, Kulkami AR and Rudzinski WE (2001) Biodegradable 
polymeric nanoparticles as drug delivery devices. / Control Rel 70:1-20. 
5. Moghimi SM, Hunter AC and Murray JC (2001) Long-circulating and target-specific 
nanoparticles: Theory to practice. Pharmacol Rev 53:283-318. 
6. Salgueiro A, Egea MA, Espina M, Vails O and Garcia ML (2004) Stability and ocular 
tolerance of cyclophosphamide-loaded nanospheres. J Microencapsul 21:213-223. 
7. Rabinow BE (2004) Nanosuspensions in drug delivery. Nat Rev Drug Discov 3: 
785-796. 
8. Moghimi SM (2002) Chemical camouflage of nanospheres with a poorly reactive surface: 
Towards development of stealth and target-specific nanocarriers. Biochim Biophys Acta 
(Mol Cell Res) 1590:131-139. 
9. Porter CJH, Moghimi SM, Ilium L and Davis SS (1992) The polyoxyethylene/ 
polyoxypropylene block co-polymer poloxamer-407 selectively redirects intravenously 
injected microspheres to sinusoidal endothelial cells of rabbit bone marrow. FEBS Lett 
305:62-66. 
10. Kreuter J, Ramge P, Petrov V, Hamm S, Gelperina SE, Engelhardt B, Alyautdin R, 
von Briesen H and Begley DJ (2003) Direct evidence that polysorbate-80-coated 
poly(butylcyanoacrylate) nanoparticles deliver drugs to the CNS via specific mechanisms 
requiring prior binding of drugs to the nanoparticles. Pharm Res 20:409-416. 
11. Huang G, Gao J, Hu Z, St. John JV, Ponder BC and Moro D (2004) Controlled drug release 
from hydrogel nanoparticle network. / Control Rel 94:303-311. 
12. Panyam J, Zhou WZ, Prabha S, Sahoo SK and Labhasetwar V (2002) Rapid endolysosomal 
escape of poly(DL-lactide-co-glycolide) nanoparticles: Implications for drug 
and gene delivery. FASEB J 16:1217-1226. 
13. St. John JV, Moro DG, Russell-Jones GJ and McDougall F (2004) Protein release from 
and cellular infiltration into hydrogel nanoparticle scaffolds. 31st Annual Meeting of the 
Controlled Release Society, Honolulu, Hawaii, June 12-16. 
14. Scholes PD, Coombes AGA, Ilium L, Davis SS, Vert M and Davies MC (1993) The 
preparation of sub-500 nm poly(lactide-co-glycolide) microspheres for site-specific drug 
delivery. / Control Rel 25:145-153. 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 39 
15. Desgouilles S, Vauthier C, Bazile D, Vacus J, Grossiord JL, Veillard M and Couvreur P 
(2003) The design of nanoparticles obtained by solvent evaporation: A comparative 
study. Langmuir 19:9504-9510. 
16. Niwa T, Takeuchi H, Hino T, Kunou N and Kawashima Y (1993) Preparations 
of biodegradable nanospheres of water-soluble and insoluble drugs with D,Llactide/
glycolide copolymer by a novel spontaneous emulsification solvent diffusion 
method and the drug release behavior. / Control Rel 25:89-98. 
17. Quintanar-Guerrero D, Allemann E, Fessi H and Doelker E (1998) Preparation techniques 
and mechanisms of formation of biodegradable nanoparticles from preformed polymers. 
Drug Dev Ind Pharm 24:1113-1128. 
18. Bilati U, Allemann E and Doelker E (2005) Development of a nanoprecipitation method 
intended for the entrapment of hydrophilic drugs into nanoparticles. Eur } Pharm Sci 24: 
67-75. 
19. Couvreur P, Kante M, Roland M, Guiot P, Bauduin P and Speiser P (1979) Polycyanoacrylate 
nanocapsules as potential lysosomotropic carriers: Preparation, morphological and 
sorptive properties. / Pharm Pharmacol 31:331-332. 
20. Lescure F, Zimmer C, Roy D and Couvreur P (1992) Optimization of polycyanoacrylate 
nanoparticle preparation: Influence of sulfur dioxide and pH on nanoparticle characteristics. 
/ Coll Interf Sci 154:77-86. 
21. De Keyser JL, Poupaert JH and Dumont P (1991) Poly(diethylmethylidenemalonate) 
nanoparticles as a potential drug carrier: Preparation, distribution and elimination after 
intravenous and peroral administration to mice. / Pharm Sci 80:67-70. 
22. Jung T, Breitenbach A and Kissel T (2000) Sulfobutylated poly(vinylalcohol)-graftedpoly(
lactide-co-glycolide) facilitate the preparation of small negatively charged 
biodegradable nanospheres for protein delivery. / Control Rel 67:157-169. 
23. Wu C, Fu J and Zhao Y (2000) Novel nanoparticles formed via self-assembly of 
poly(ethylene glycol-b-sebacic anhydride) and their degradation in water. Macromolecules 
33:9040-9043. 
24. Broz P, Benito SM, Saw CL, Burger P, Heider H, Pfisterer M, Marsch S, Meier W and 
Hunziker P (2005) Cell targeting by a generic receptor-targeted polymer nanocontainer 
platform. / Control Rel 102:475^88. 
25. Boudad H, Legrand P, Lebas G, Cheron M, Duchene D and Ponchel GG (2001) Combined 
hydroxylpropyl-beta-cyclodextrin and poly(alkylcyanoacrylate) nanoparticles intended 
for oral administration of saquinavir. Int} Pharm 218:113-124. 
26. Labib A, Lenaerts V, Chouinard F, Leroux JC, Ouellet R and van Lier JE (1991) 
Biodegradable nanospheres containing phthalocyanines and naphthalocyanines for targeted 
photodynamic tumor therapy. Pharm Res 8:1027-1031. 
27. Jeong YI, Cho CS, Kim SH, Ko KS, Kim SI, Shim YH and Nah JW (2001) Preparation 
of poly(DL-lactide-co-glycolide) nanoparticles without surfactant. / Appl Polym Sci 80: 
2228-2236. 
28. Allemann E, Leroux JC, Gurnay R and Doelker E (1993) In vitro extended-release 
properties of drug-loaded poly(D,L-lactic) acid nanoparticles produced by a salting-out 
procedure. Pharm Res 10:1732-1737. 
40 Moghimi et al. 
29. Randolph TW, Randolph AD, Mebes M and Yeung S (1993) Submicron-sized biodegrad-. 
able particles of poly(L-lactic acid) via the gas antisolvent spray precipitation process. 
Biotechnol Prog 9:429-435. 
30. Tokumitsu H, Ichikawa H and Fuukumori Y (1999) Chitosan-gadopenteic acid complex 
nanoparticles for gadolinium neutron-capture therapy of cancer: Preparation by novel 
emulsion-droplet coalscence technique and charaterization. Pharm Res 16:1830-1835. 
31. Prokop A, Kozlov E, Newman GW and Newman MJ (2002) Water-based nanoparticulate 
polymeric system for protein delivery: Permeability control and vaccine application. 
Biotechnol Bioeng 78:459^166. 
32. Calvo P, Remunan-Lopez C, Vila-Jato JL and Alonso MJ (1997) Chitosan and chitosan/
ethylene oxide-propylene oxide block copolymer nanoparticles as novel carriers 
for proteins and vaccines. Pharm Res 14:1431-1436. 
33. Liu H, Finn N and Yates MZ (2005) Encapsulation and sustained release of a model drug, 
indomethacin, using C02-based microencapsulation. Langmuir 21:379-385. 
34. Panyam J, Williams D, Dash A, Leslie-Pelecky D and Labhasetwar V (2004) Solid-state 
solubility influences encapsulation and release of hydrophobic drugs from PLGA/PLA 
nanoparticles. f Pharm Sci 93:1804-1814. 
35. Polakovic M, Gorner T, Gref R and Dellacherie E (1999) Lidocaine loaded biodegradable 
nanospheres. II. Modeling of drug release. / Control Rel 60:169-177. 
36. Tamber H, Johansen P, Merkle HP and Gander B (2005) Formulation aspects of 
biodegradable polymeric microspheres for antigen delivery. Adv Drug Del Rev 57: 
357-376. 
37. Stella B, Arpicco S, Peracchia MT, Desmaele D, Hoebeke J, Renoir M, D'Angelo J, Cattel L 
and Couvreur P (2000) Design of folic acid-conjugated nanoparticles for drug targeting. 
/ Pharm Sci 89:1452-1464. 
38. Moghimi SM, Porter CJH, Muir IS, Ilium L and Davis SS (1991) Non-phagocytic uptake 
of intravenously injected microspheres in rat spleen: Influence of particle size and 
hydrophilic coating. Biochem Biophys Res Commun 177:861-866. 
39. Moghimi SM, Hedeman H, Ilium L and Davis SS (1993) Effect of splenic congestion 
associated with haemolytic anaemia on filtration of "spleen-homing" microspheres. Clin 
Sci 84:605-609. 
40. Moghimi SM, Hedeman H, Christy NM, Ilium L and Davis SS (1993) Enhanced hepatic 
clearance of intravenously administered sterically stabilized microspheres in zymosanstimulated 
rats. / Leukoc Biol 54:513-517. 
41. Laverman P, Carstens MG, Storm G and Moghimi SM (2001) Recognition and clearance 
of methoxypoly(ethyleneglycol) 2000-grafted liposomes by macrophages with enhanced 
phagocytic capacity. Implications in experimental and clinical oncology. Biochim Biophys 
Acta (General Subjects) 1526:227-229. 
42. Calvo P, Gouritin B, Villarroya H, Eclancher F, Giannavola C, Klein C, Andreux JP 
and Couvreur P (2002) Quantification and localization of PEGylated polycyanoacrylate 
nanoparticles in brain and spinal cord during experimental allergic encephalomyelitis 
in the rat. Eur } Neurosci 15:1317-1326. 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 41 
43. Song C, Labhasetwar V, Cui X, Underwood T and Levy RJ (1998) Arterial uptake of 
biodegradable nanoparticles for intravascular local drug delivery: Results with an acute 
dog model. / Control Rel 54:201-211. 
44. Moghimi SM and Bonnemain B (1999) Subcutaneous and intravenous delivery of diagnostic 
agents to the lymphatic system: Applications in lymphoscintigraphy and indirect 
lymphography. Adv Drug Del Rev 37:295-312. 
45. Hawley AE, Ilium L and Davis SS (1997) Lymph node localisation of biodegradable 
nanospheres surface modified with poloxamer and poloxamine block co-polymers. FEBS 
Lett 400:319-323. 
46. Moghimi SM (2003) Modulation of lymphatic distribution of subcutaneously injected 
poloxamer 407-coated nanospheres: The effect of the ethylene oxide chain configuration. 
FEBS Lett 540:241-244. 
47. Moghimi SM and Rajabi-Siahboomi AR (1996) Advanced colloid-based systems for efficient 
delivery of drugs and diagnostic agents to the lymphatic tissues. Prog Biophys Mol 
Biol 65:221-249. 
48. Jiang W, Gupta RK, Deshpande MC and Schwendeman SP (2005) Biodegradable poly 
(lactic-co-glycolic acid) microparticles for injectable delivery of vaccine antigens. Adv 
Drug Del Rev 57:391^10. 
49. Simecka JW (1998) Mucosal immunity of the gastrointestinal tract and oral tolerance. 
Adv Drug Del Rev 34:235-259. 
50. Ermak TH and Giannasca PJ (1988) Microparticle targeting to M cells. Adv Drug Deliv 
Rev 34:261-283. 
51. O'Hagan DT and Valiante NM (2003) Recent advances in the discovery and delivery of 
vaccine adjuvants. Nat Rev Drug Discov 2:727-735. 
52. O'Hagan DT, Singh M and Ulmer JB (2004) Microparticles for delivery of DNA vaccines. 
Immunol Rev 199:191-200. 
53. van der Lubben IM, Kersten G, Fretz MM, Beuvery C, Verhoef JC and Junginger HE 
(2003) Chitosan microparticles for mucosal vaccination against diphteria: Oral and nasal 
efficacy studies in mice. Vaccine 21:1400-1408. 
54. Takeuchi H, Yamamoto H and Kawashima YY (2001) Mucoadhesive nanoparticulate 
systems for peptide drug delivery. Adv Drug Del Rev 47:39-54. 
55. Carino GP, Jacob JS and Mathiowitz E (2000) Nanosphere based oral insulin delivery. 
/ Control Rel 65:261-269. 
56. Arbos P, Campanero MA, Arangoa MA, Renedo MJ and Irache JM (2003) Influence of 
the surface characteristics of PVM/MA nanoparticles on their bioadhesive properties. 
/ Control Rel 89:19-30. 
57. Tobio M, Sanchez A, Vila A, Soriano I, Evora C, Vila-Jato JL and Alonso MJ (2000) The 
role of PEG on the stability in digestive fluids and in vivo fate of PEG-PLA nanoparticles 
following oral administration. Coll Surf (B-Biointerface) 18:315-323. 
58. Fresta M, Fontana C, Bucolo G, Cavallaro G, Giammona G and Puglisi G (2001) Ocular 
tolerability and in vivo bioavailability of poly(ethylene glycol) (PEG)-coated polyethyl- 
2-cyanoacrylate nanospheres-encapsulated acyclovir. / Pharm Sci 90:288-297. 
42 Moghimi et al. 
59. de Kozak Y, Andrieux K, Villarroya H, Klein C, Thillaye-Goldenberg B, Naud MC, 
Garcia E and Couvreur P (2004) Intraocular injection of tamoxifen-loaded nanoparticles: 
Anew treatment of experimental autoimmune uveoretinitis. Eur J Immunol 34:3702-3712. 
60. Giannavola C, Bucolo C, Maltese A, Paolino D, Vandelli MA, Puglisi G, Lee VHL and 
Fresta M (2003) Influence of preparation conditions on acyclovir-loaded poly-d,l-lactic 
acid nanospheres and effect of PEG coating on occular drug bioavailability. Pharm Res 
20:584-590. 
61. Bucolo C, Maltese A, Maugeri F, Busa B, Puglisi G and Pignatello R (2004) Eudragit 
RL100 nanoparticle system for the opthalmic delivery of cloricromene. / Pharm Pharmacol 
56:841-846. 
62. Vila A, Gill H, Mccallion O and Alonso MJ (2004) Transport of PLA-PEG particles across 
the nasal mucosa: Effect of particle size and PEG coating density. / Control Rel 98:231-244. 
63. Moghimi SM, Hunter AC, Murray JC and Szewczyk A (2004) Cellular distribution of 
nonionic micelles. Science 303:626-627. 
64. Hunter AC and Moghimi SM (2002) Therapeutic synthetic polymers: A game of Russian 
roulette? Drug Discov Today 7:998-1001. 
65. Moghimi SM, Symonds P, Murray JC, Hunter AC, Debska G and Szewczyk A 
(2005) A two-stage poly(ethylenimine)-mediated cytotoxicity: Implications for genetransfer 
/therapy. Mol Ther 11:990-995. 
66. Gbadamosi JK, Hunter AC and Moghimi SM (2002) PEGylation of microspheres generates 
a heterogeneous population of particles with differential surface characteristics and 
biological performance. FEBS Lett 532:338-344. 
4 
Genetic Vaccines: A Role for Liposomes 
Gregory Gregoriadis, Andrew Bacon, 
Brenda McCormack and Peter Laing 
1. Introduction 
Prevention of microbial infections by the use of vaccines is a preferred alternative 
to treatment. Vaccines have been applied successfully, for example, in the eradication 
of smallpox as well as against tetanus, diphtheria, whooping cough, polio and 
measles, thus preventing millions of deaths each year. However, vaccines made 
of attenuated organisms, which mimick natural infections usually without the disease, 
can be potentially unsafe. For instance, there is a risk of reversion during replication 
of live viruses or even mutation to a more pathogenic state. Furthermore, 
with immunocompromised individuals, some of the attenuated viruses may still 
provoke disease. On the other hand, with killed virus vaccines, their extracellular 
localization and subsequent phagocytosis by professional antigen presenting cells 
(APC) or antigen-specific B cells, lead to MHC-II class restricted presentation and to 
T helper cell and humoural immunity. However, they do not elicit significant cytotoxic 
T cell (CTL) responses. Moreover, subunit vaccines produced from biological 
fluids may not be entirely free of infectious agents. Even with subunit and peptide 
vaccines produced recombinantly or synthetically (and thus considered safe), 
immune responses are weak and often not of the appropriate kind. The great variety 
of immunological adjuvants1'2 that are now available go a long way in rendering 
subunit and peptide vaccines stronger and more efficient. However, more than seventy 
years after the introduction of aluminium salts as an adjuvant, only two other 
adjuvants, liposomes3 and MF59,1 have been approved for use in humans.4 Thus, 
43 
44 Gregoriadis et al. 
inspite of considerable progress, the road to the ideal vaccine appears as elusive as 
ever, until recently. 
Recent developments have led to a novel and exciting concept, namely de novo 
production of the required vaccine antigen by the host's cells in vivo, which promises 
to revolutionize vaccination especially where vaccines are either ineffective or 
unavailable. The concept entails the direct injection of antigen-encoding plasmid 
DNA which, on uptake by cells, localizes to some extent into the nucleus where 
it transfects the cells episomally. The produced antigen is recognized as foreign 
by the host and is thus subjected to pathways similar to those observed for antigens 
of internalized viruses (but without their disadvantages), leading to protective 
humoural and cell mediated immunity.5-9 A series of publications since 1992 first 
established the ability of plasmid DNA to induce an immune (antibody) response 
to the encoded foreign protein10; in experiments with DNA encoding influenza 
nucleoprotein, immunity was both humoural and cell-mediated, and also protective 
in mice challenged with the virus.11,12 This was the first demonstration of an 
experimental DNA vaccine. Another observation was the induction of humoural 
and cell-mediated immunity against HIV-1 using plasmids encoding the HIV rev 
and env proteins.13 Similar results were obtained with a gene for the hepatitis B 
surface antigen (HBsAg).14 DNA immunization was also found to apply in cancer 
treatment. For instance, injection of plasmids encoding tumor antigens promoted 
immune responses15,16 which were protective in an animal model.6 The concept 
of DNA immunization has now been adopted by vaccinologists worldwide using 
an ever increasing number of plasmids encoding immunogens from bacterial, viral 
and parasitic pathogens, and a variety of tumors.8,9 In many of these studies, genetic 
immunization has led to the protection of animals from infection.5-9 A number of 
clinical trials for the therapy of, or prophylaxis against, a variety of infections are 
in progress.8,9 
2. The DNA Vaccine 
A plasmid DNA vaccine is usually6 supercoiled and consists of the gene encoding 
the vaccine antigen (the section of the target pathogen which elicits protective 
immunity), a promoter sequence which is often derived from cytomegalovirus 
(CMV) or Rous sarcoma virus (RSV)), an mRNA stability polyadenylation region 
at the 3' end of the insert, and the plasminogen activator gene which controls the 
secretion of the recombinant product. In addition, there are an origin of replication 
for the amplification of the plasmid in bacteria, and a gene for antibiotic resistance 
to select the transformed bacteria. 
Immunization procedures with DNA vaccines are carried out by the intramuscular 
and, to a lesser extent, the intraepidermal route. Other routes include the 
Genetic Vaccines: A Role for Liposomes 45 
oral, nasal, vaginal, intravenous, intraperitoneal and subcutaneous routes.8,9 Intramuscular 
injection of DNA vaccines leads to such types of immunity as CTL.5'9,11'12 
This was unexpected because antigen presentation requires the function of professional 
APC.17 However, myocytes which were shown5 to take up the plasmid only 
to a small extent and with only a fraction of cells participating in the uptake, are 
not professional APCs. Although myocytes carry MHC class I molecules and can 
present endogenously produced viral peptides to the CD8+ cells to induce CTLs, 
they do so inefficiently18 as they lack vital costimulatory molecules (e.g. the B7-1 
molecule). It is thus difficult to accept that antigen presentation, leading to a CTL 
response, occurs via myocytes. Instead, it was reported18 that CTL responses occur 
as a result of the transfer of antigenic material between the myocytes and professional 
APC to some extent. In parallel, it could also be that plasmid secreted by 
the myocytes or as such, is taken up directly by APC infiltrating the injected site. 
Such APC would include dendritic cells which will express and present peptides 
to CD8+ cells following transport to the lymph nodes or spleen. On the other hand, 
CD4+ cells may be activated by APCs via MHC class II presentation of antigen 
secreted by the myocytes (or released from them after their destruction via a Tc 
response) and captured by the cells. Such events would lead to both cellular (Th 1) 
and humoural (Th 2) immunity. Indeed, it has been shown6 that dendritic cells are 
the essential APC involved in immune responses elicited by intramuscularly given 
DNA vaccines. 
3. DNA Vaccination via Liposomes 
Vaccination with naked DNA by the intramuscular route is dependent on the ability 
of myocytes to take up the plasmid. However, some of the DNA may also be 
engulfed by APC infiltrating the site of injection, or in the lymph nodes following 
migration of the DNA to the lymphatics. The extent of DNA degradation by 
extracellular deoxyribonucleases is unknown, but depending on the time of its residence 
interstitially, degradation could be considerable. Therefore, approaches that 
protect DNA from the extracellular nucleases and promote DNA uptake by cells 
more efficiently, or target it to APC, should contribute to the optimal design of DNA 
vaccines. 
It has been suggested19 that as APC are a preferred alternative to muscle cells 
for DNA vaccine uptake and expression, liposomes (known3 to be taken up avidly 
by APC infiltrating the site of injection or in the lymphatics, an event that has 
been implicated3 in their immunoadjuvant activity) would be a suitable means 
of delivery of entrapped DNA to such cells. Liposomes would also protect20 their 
DNA content from deoxyrubonuclease attack. Moreover, the structural versatility21 
of the system would ensure that its tranfection efficiency is further improved 
46 Cregoriadis ct al. 
by the judicial choice of its structural characteristics or by the co-entrapment of 
cytokine genes, other adjuvants (e.g. immunostimulatory sequences), or indeed 
protein antigens (see later) together with the plasmid vaccine. As a number of 
injectable liposome-based drug formulations, including vaccines against hepatitis A 
and influenza, have been already licensed for clinical use,21 acceptance of the system 
clinically would be less problematic than with other systems that are still at an 
experimental stage. 
3.1. Procedure for the entrapment of plasmid DNA into liposomes 
A variety8,22,23 of plasmid DNAs have been quantitatively entrapped into liposomes 
by a mild dehydration-rehydration procedure.20'22,23 The procedure (Fig. 1) 
consists of mixing preformed small unilamellar vesicles (SUV) with a solution 
of the DNA destined for entrapment, freeze-drying of the mixture, followed by 
controlled rehydration of the formed powder, and centrifugation to remove nonentrapped 
material. Formed liposomes are multilamellar.20 However, when an 
appropriate amount of sucrose is added to the SUV and DNA mixture prior 
to dehydration,24 the resulting liposomes are much smaller (about 100-160 nm 
in diameter). As expected, DNA incorporation values8'23-26 were higher (up to 
90% of the amount used) when a cationic lipid was present in the bilayers. No 
apparent relationship was observed between amount of DNA used (10-500//g) 
and the values of incorporation for the compositions and lipid mass used.8,23,26 
The possibility that DNA was not entrapped within the bilayers of cationic liposomes, 
but was rather complexed with their surface (as suggested by the high 
Fig. 1. Entrapment of DNA and/or protein into cationic liposomes. The procedure entails 
mixing up empty SUV with the solute(s) destined for entrapment and subsequent dehydration. 
On rehydration, most of the solute(s) is recovered entrapped within the generated 
multilamellar liposomes. 
Genetic Vaccines: A Role for Liposomes 47 
Naked DNA 
IMBXM™ (DNA) § 
m 
"Complexed** DNA 
Naked DNA 
taraXeB™ (DNA) g 
"Complexed" DNA 
4* 
W 
o 
Fig. 2. Gel electrophoresis of a mixture of cationic SUV and pRc/CMV HBS before (complexed 
DNA) and after (entrapped DNA) dehydration-rehydration of the mixture. 
"incorporation" values obtained on mixing)20 was examined by treating liposomeentrapped 
and liposome-complexed DNA with deoxyribonuclease. Substantially, 
more liposome-entrapped DNA remained intact than when it was complexed,20 
presumably because of the inability of the enzyme to reach its substrate in the former 
case. The significant resistance of complexed DNA (despite its accessibility) 
to the enzyme could be attributed to its condensed state.25 Additional evidence 
that the DNA was entrapped within liposomes was obtained by gel electrophoresis 
of a mixture of cationic SUV and plasmid DNA before (complexed DNA) and 
after dehydration-rehydration of the mixture (entrapped DNA). When the anionic 
sodium dodecylsulphate (SDS) was incorporated in the gel, complexed DNA was 
dissociated from the SUV, presumably because of ionic competition for the cationic 
charges. As expected, "entrapped" DNAretained its association with the liposomes, 
suggesting its unavailability to the competing SDS anions26 (Fig. 2). 
3.2. DNA immunization studies 
Previously,20 liposome-entrapped plasmid found to transfect cells in vitro regardless 
of the vesicle surface charge was tested in immunization experiments,19,27 using 
a plasmid (pRc/CMV HBS) encoding the S region of the hepatitis B surface antigen 
(HBsAg; subtype ayw). Mice (Balb/c) that are repeatedly injected intramuscularly 
with 5 or 10/ig plasmid entrapped in cationic liposomes, exhibited at all 
times much greater (up to 100-fold) antibody (IgGi) responses (Fig. 3) against the 
48 Gregoriadis et al. 
© 
I 
O 
| 
E 
c 5 a. I 
a. 
8 I 
D 
JBfcei_ 
26 34 44 
Days after first injection 
Fig. 3. Immune responses in mice injected with naked, or liposome-entrapped pRc/CMV 
HBS. Balb/c mice were injected intramuscularly on days 0, 10, 20, 27 and 37 with 5 /xg of 
DNA entrapped in cationic liposomes composed of PC, DOPE and DOTAP (A), DC-Chol 
(B) or SA (C) (molar ratios 1:0.5:0.25), or in the naked form (D). Animals were bled 7, 15, 
26, 34 and 44 days after the first injection and sera tested by ELISA for IgGT (black bars), 
IgG2a (white bars) or IgG2b (grey bars) responses against the encoded hepatitis B surface 
antigen (HBsAg; S region, ayw subtype). Values are means ±SD of log10 of reciprocal end 
point serum dilutions required for OD to reach readings of about 0.2. Sera from untreated 
mice gave log10 values of less than 2.0. IgGj responses were mounted by all mice injected 
with liposomal DNA but became measurable only at 26 days. Differences in log10 values 
(all IgG subclasses at all time intervals) in mice immunized with liposomal DNA and mice 
immunized with naked DNA were statistically significant (P < 0.0001-0.002). (Reproduced 
with permission from Ref. 19.) 
Genetic Vaccines: A Role for Liposomes 49 
encoded antigen than animals immunized with the naked plasmid. Values of other 
subclasses (IgG2a and IgG2b) were also greater (up to 10-fold) (Fig. 3). Moreover, 
IgGj responses for the liposome-entrapped plasmid DNA were higher (up to 10- 
fold) than those obtained with DNA complexed with similar cationic liposomes.19 
This was also true for IFN-y and IL-4 levels in the spleens of immunized mice.19 
In other experiments,8 the effect of the route of injection of the pRc/CMV HBS 
plasmid was examined with respect to both humoural and cell-mediated immunity, 
using Balb/c mice and an outbred mouse strain (T.O.). Results8 comparing 
responses for liposome-entrapped and naked plasmid DNA showed greater antibody 
(IgGi) responses for the entrapped DNA, not only by the intramuscular route, 
but also the subcutaneous and the intravenous routes. As there were no significant 
differences in the titers between the two strains,8 it was concluded that immunization 
with liposomal pRc/CMV HBS is not MHC restricted. Results obtained 
on the testing of IFN-y and IL-4 in the spleens (not shown) exhibited a similar 
pattern. 
Involvement of muscle cells in the mechanism by which liposomes promote 
greater immune responses to the encoded antigen than seen with the naked plasmid, 
is rather unlikely. Although, cationic liposomes could in theory bind to and 
be taken up by the negatively charged myocytes, the negatively charged proteins 
present in the interstitial fluid would neutralize21 the cationic liposomal surface 
and thus interfere with such binding. In addition, vesicle size (about 600-700 nm 
average diameter; Ref. 26) would render access to the cells difficult, if not impossible. 
It is therefore more likely that cationic liposomes are endocytosed by APC, 
including dendritic cells, in the lymphatics where liposomes are expected to end 
up.28 Uptake of liposomal plasmid DNA is supported in studies where mice were 
injected intramuscularly or subcutaneously with liposomes entrapping the plasmid 
(pCMV- EFGP), encoding the enhanced fluorescent green protein or with the 
naked plasmid. Fluorescence microscopy of sections of the lymph nodes draining 
the injected site revealed (Fig. 4) much more green fluorescence when the plasmid 
was administered in the liposomal form.27 It appears8'19 that the key ingredient of 
the DNA-containing liposomes as used in Fig. 3, contributing to enhanced immune 
responses, is the cationic lipid. The mechanism by which liposomal DNAreaches the 
nucleus for episomal transfection is poorly understood. It is conceivable, however, 
that some of the endocytosed liposomal DNA escapes the endocytic vacuoles prior 
to their fusion with lysosomes (in a way similar to that proposed29 for vesicle-DNA 
complexes) to enter the cytosol for eventual episomal transfection and presentation 
of the encoded antigen. It is perhaps at this stage of intracellular trafficking of DNA, 
spanning its putative escape from endosomes and access to the nucleus, that the 
cationic lipid, possibly together within the fusogenic phosphatidylethanolamine 
(PE) component, plays a significant role. 
50 Gregoriadis et al. 
Kttmam w»* ii'm<|»* 
IWt 
•ff!f b ((lit 
tyfent 
Fig. 4. Fluorescence images of muscle and lymph node sections from mice injected intramuscularly 
with 10/xg liposome-entrapped or naked pCMVEGFP and killed 48h later. 
Sections from untreated animals were used as controls. (Reproduced with permission from 
Ref. 27.) 
3.3. Induction of a cytotoxic T lymphocyte (CTL) response 
by liposome-entrapped plasmid DNA 
Immunization studies with liposome-entrapped DNA vaccines were expanded30 
to include the cytotoxic T lymphocyte (CTL) component of the immune response. 
This was measured by the specific killing of syngeneic target cells pulsed with a 
recognized CTL epitope peptide derived from the antigen tested. To that end, the 
type and degree of immune response induced following subcutaneous injection of 
DNA in cationic liposomes was monitored and compared with that obtained with 
DNA alone injected by the same route. 6-8 week old, female C57/BL6 (H-2d) mice 
were injected subcutaneously with one or two doses of 2.5 or 10 ^g ovalbumin 
(OVA)-encoding plasmid DNA (pCI-OVA), either alone or entrapped in liposomes. 
Animals immunized subcutaneously with 100 /xg of OVA protein complexed with 
1 /xg of cholera toxin (CT) served as a positive control. Blood samples and spleens 
were collected from all animals one week after the last injection and tested for 
anti-OVA total IgG (serum), CTL activity and cytokine release (spleen). After a 
single dose of antigen, only animals immunized with either protein or 10/xg of 
liposomal DNA showed significant anti-OVA antibody titres by ELBA. After two 
doses of antigen, only animals immunized with either protein or liposomal DNA 
(both 2.5 and 10 ttg DNA) showed significant levels of seroconversion and serum 
antibody titres against OVA by ELBA.30 Similarly, no anti-OVA CTL activity was 
detected in animals immunized with DNA alone. However, animals immunized 
with two doses of 10 /xg of liposomal DNA displayed a CTL response higher (60% 
cell killing vs 50%) than that obtained in the positive control group immunized 
Genetic Vaccines: A Role for Liposomes 51 
with OVA protein and adjuvant (CT).30 Thus, delivery of a small dose of liposomal 
plasmid DNA subcutaneously, a route of immunization not normally inducing 
significant plasmid DNA mediated immune activation,9 results in a strong antigenspecific 
cellular response which is greater than that achieved by higher doses of a 
conventional protein antigen together with a powerful adjuvant (CT). 
4. The Co-delivery Concept 
Proteins that are synthesized within a cell (e.g. from plasmid DNA having a 
mammalian-active promoter) are continuously sampled as peptides by the 
proteosome / class-I MHC antigen presenting pathway. Conversely, proteins that are 
acquired exogenously by antigen-presenting cells are sampled in an analogous way 
by the endosomal/MHC-class-II pathway. It follows that the delivery of both protein 
and plasmid-DNA-encoded forms of a protein antigen to the same individual 
antigen-presenting cell would result in the simultaneous presentation of the antigen 
via both class-I and class-II pathways, thereby providing an opportunity for synergy 
in the resulting immune response to the antigen. Several appropriate liposomal 
formulations were designed to test the "co-delivery" hypothesis, exploiting the 
advantages of the dehydration-rehydration liposome technology that entraps both 
DNA and protein immunogens efficiently. The formulations, described in Table 1, 
comprise various test and control permutations of plasmid DNA and protein, either 
free or entrapped (together or separately) in the liposomal vehicle. 
Immunization with DNA encoding the influenza haemagglutinin protein 
has been explored previously with naked31 or liposomally formulated DNA.32 
Although immune responses elicited by DNA alone were adequate to achieve protective 
efficacy against influenza virus challenge in preclinical studies, only weak 
anti-HA antibody responses were elicited.31 The present "co-delivery" concept was 
designed to rectify this deficiency of DNA-based influenza vaccines. In a series of 
experiments, plasmid DNA encoding the haemagglutinin (HA) antigen [referred to 
in Table 1 and Fig. 5 as DNA(ha)] of the influenza virus (A/Sichuan/87 or A/PR/8 
strains) was co-entrapped with the corresponding whole inactivated virus (referred 
to as HA) within the same liposomes using the dehydration-rehydration method 
(for details on lipid composition and method see Refs. 26 and 27). A variety of control 
preparations including liposomes co-entrapping irrelevant DNA (i.e. plasmid 
DNA encoding ovalbumin) with HA or irrelevant protein (i.e. ova) with DNA (ha), 
entrapping DNA(ha) or HA alone, a mixture of the latter two preparations, and 
a mixture of the naked DNA(ha) and HA were used to immunize mice. Results 
shown in Fig. 5 demonstrate that the "co-delivery" hypothesis formulation (comprising 
both HA and its corresponding DNA in the same liposomes), elicited a 
greater response than all other formulations at each time point in the series, and it 
52 Gregoriadis et al. 
Table 1 Liposomal formulations of DNA and protein used in immunization experiments. 
Sample 
1.1 
2.1 
3.1 
4.1 
5.1 
6.1 
7 
8 
9 
10 
11 
12 
Dose (/ig/animal (0.2 ml S/O) 
Formulation 
Liposomes 
(co-delivery) 
Liposomes 
(co-delivery) 
Liposomes 
(co-delivery) 
Liposomes 
Liposomes 
Liposomes (samples 4.1 & 5.1) 
DNA and protein (mixed) 
DNA and protein (mixed) 
DNA and protein (mixed) 
DNA alone 
Protein alone 
Control (PBS) 
DNA 
ha (10) 
ova (11) 
ha (10) 
ha (10) 
Nil 
ha (10) 
ha (10) 
ova (11) 
ha (10) 
ha (10) 
Nil 
Nil 
Protein 
HA (0.6) 
HA (0.6) 
OVA (0.76) 
Nil 
HA (0.6) 
HA (0.6) 
OVA (0.76) 
HA (0.6) 
HA (0.6) 
Nil 
HA (10) 
Nil 
Plasmid DNA encoding the HA antigen [DNA(ha)] and the HA antigen (HA) were entrapped in liposomes 
either together (co-entrapped; sample 1.1) or separately in different formulations (sample 6.1) 
mixed before injection. In some formulations, DNA(ha) and HA were entrapped alone (samples 4.1 and 
5.1 respectively). In others, ovalbumin (OVA) and plasmid DNA encoding ha fDNA(ha)] (sample 7) 
or HA and plasmid DNA encoding OVA (sample 8) were entrapped separately and then mixed. Mice 
were injected subcutaneously on days 0 and 28 and blood samples analyzed by ELISA for Ig responses. 
1OD0O - 
1000 
100- 
A/Sichuan/87 
Ig response 
-p=o.oca 
;r***OM burton) 
DNA (10 (ig) / Protein (0.6 ng) 
- • - Up(DNA(HA)/HA) 
- * - Lip(ONA{OVA)/HA) 
- * - Up(DNA(HA)/OVA) 
- • - Up (DNA (HA)/no protein) 
- • - Up(noDNA/HA) 
- * - Up(DNA(H;)) + Up(HA) 
•••••• DNA {HA ) • OVA 
• DNA (OVA)* HA 
• DNA(HA)+HA 
- * - DNA ( n » ) no protein 
- * - HP (protein alone) 
• control (negative J 
20 A 3 0 
boost 
40 50 
Day post 1st dose 
Fig. 5. Serum Ig endpoint titres in Balb/c mice immunized on days 0 and 28 with DNA 
and/or antigen formulations as described in Table 1 a nd bled at time intervals. 
Genetic Vaccines: A Role for Liposomes 53 
is by far the strongest response after a single dose. Notably, the formulation "Lip 
(OVA/ha)", which is a control for the CpG adjuvant effect of plasmid DNA,33 gave 
a response which was much lower than that of "co-delivery" with the appropriate 
homologous pair of HA DNA and protein. Likewise, Lip (HA/ova) (an inappropriate 
pairing according to the hypothesis), gave a markedly weaker response. Figure 5 
also demonstrates that separately entrapped HA DNA and protein (in neighbouring 
vesicles) gave rise to an inferior response, supporting the hypothesis that delivery 
of both payloads to the same cell (which is best achieved by co-entrapment 
in the same liposome) is important in achieving the optimal antibody response. It 
is also remarkable that, inspite the modest DNA dose (10 /xg) and small number 
(2) of immunizations used, several formulations completely failed to generate an 
anti-HA response. These included HA DNA alone, and liposomally entrapped HA 
DNA. These findings serve to emphasize the striking degree of superiority of "codelivery" 
over previous methods of DNA-based immunization against influenza 
virus. 
In conclusion, the present studies demonstrate that very small doses of protein 
as an additive in DNA immunization can dramatically improve the antibody 
response to the target protein, provided that the protein and DNA are homologous 
to one-another (i.e. that the DNA can express the protein), and that the payloads 
are delivered in the same individual liposomal vehicle. The simplest hypothesis 
to explain our observation is that the synergy observed between the appropriately 
delivered "homologous pair" of protein and DNA involves delivery of both 
payloads to the same antigen-presenting cell. The application of the co-delievery 
concept to alternative delivery systems, e.g. niosomes, dendimers, PLA/PLGA, chitosans, 
alginates and other microparticles awaits investigation. It is anticipated that 
the "co-delivery" approach will lead to better DNA-based vaccines for prophylactic 
and therapeutic use, particularly where vaccines require the elicitation of antibody 
responses (e.g. influenza vaccines). 
References 
1. Powel MF and Newman MJ (eds.) (1995) Vaccine Design: The Subunit and Adjuvant 
Approach. Plenum Press: New York. 
2. Gregoriadis G, McCormack B, Allison AC and Poste G (eds.) (1993) New Generation 
Vaccines: The Role of Basic Immunology. Plenum Press: New York. 
3. Gregoriadis G (1990) Immunological adjuvants: A role for liposomes. Immunol Today. 
11:89-97. 
4. Gluck R, Mischler R, Brantschen S, Just M, Althans B and Cryz SJ, Jr (1992) Immunopotentiating 
reconstituted influenza virome vaccine delivery system for immunization 
against hepatitis A. / Clin Invest 90:2491-2495. 
54 Gregoriadis et al. 
5. Davis HL, Whalen RG and Demeneix BA (1993) Direct gene transfer in skeletal muscle 
in vivo: Factors influencing efficiency of transfer and stability of expression. Hum Gene 
Ther 4:151-156. 
6. Manickan E, Karem KL and Rouse BT (1997) DNA vaccines — A modern gimmick or a 
boon to vaccinology? Crit Rev Immunol 17:139-154. 
7. Chattergoon M, Boyer J and Weiner DB (1997) Genetic immunization: A new era in 
vaccines and immune therapeutics. FASEB 11:754-763. 
8. Gregoriadis G (1998) Genetic vaccines: Strategies for optimization. Pharm Res 15:661-670. 
9. Lewis PJ and Babiuk LA (1999) DNA vaccines: A review. Adv Virus Res 54:129-188. 
10. Tang DC, Devit M and Johnston SA (1992) Genetic immunization is a simple method for 
eliciting an immune response. Nature 356:152-154. 
11. Ulmer JB, Donnelly J, Parker SE, et al. (1993) Heterologous protection against influenza 
by injection of DNA encoding a viral protein. Science 259:1745-1749. 
12. Fynan EF> Webster RG, Fuller DH and Haynes JR (1993) DNA vaccines: Protective immunizations 
by parenteral, mucosal and gene-gun inoculations. Proc Natl Acad Sci USA 
90:11478-11482. 
13. Wang B, Ugen K, Srikantan V, et al. (1993) Gene inoculation generates immune responses 
against HIV-I. Proc Natl Acad Sci USA 90:4156^160. 
14. Davis HL, Michel ML, Mancini M, Schleef M and Whalen RG (1994) Direct gene transfer 
in skeletal muscle: Plasmid DNA based immunization against the hepatitis B virus 
surface antigen. Vaccine 12:1503-1509. 
15. Conry R, LoBuglio A, Loechel F, et al. (1995) A carcinoembryonic antigen polynucleotide 
vaccine for human clinical use. Cancer Gene Ther 2:33-38. 
16. Bright RK, Beames B, Shearer MH and Kennedy RC (1996) Protection against lethal 
tumor challenge with SV40-transformed cells by the direct injection of DNA encoding 
SV-40 large tumor antigen. Cancer Res 56:1126-1130. 
17. Matzinger P (1994) Tolerance, danger and the extended family. Annu Rev Immunol 12: 
991-1045. 
18. Spier E (1996) Meeting Report: International meeting on the nucleic acid vaccines for 
the prevention of infectious disease and regulatory nuclear acid (DNA) vaccines. Vaccine 
14:1285-1288. 
19. Gregoriadis G, Saffie R and de Souza B (1997) Liposome-mediated DNA vaccination. 
FEES Lett 402:107-110. 
20. Gregoriadis G, Saffie R and Hart SL (1996) High yield incorporation of plasmid DNA 
within liposomes: Effect on DNA integrity and transfection efficiency. / Drug Targ 3: 
469-475. 
21. Gregoriadis G (1995) Engineering targeted liposomes: Progress and problems. Trends 
Biotechnol 13:527-537. 
22. Gregoriadis G, McCormack B, Obrenovic M and Perrie Y (1999) Entrapment of plasmid 
DNA vaccines into liposomes by dehydration/rehydration, in Lowrie DB and Whalen R. 
(eds.) Methods in Molecular Medicine, DNA Vaccines: Methods and Protocols. Humana Press 
Inc.: Totowa, NJ. pp. 305-312. 
Genetic Vaccines: A Role for Liposomes 55 
23. Gregoriadis G, McCormack B, Obrenovic M, Saffie R, Zadi B and Perrie Y (1999) Liposomes 
as immunological adjuvants and vaccine carriers. Methods 19:156-162. 
24. Zadi B and Gregoriadis G (2000) A novel method for high-yield entrapment of solutes 
into small liposomes. J Lipos Res 10:73-80. 
25. Feigner PL and Rhodes G (1991) Gene therapeutics. Nature 349:351-352. 
26. Perrie Y and Gregoriadis G (2000) Liposome-entrapped plasmid DNA: Characterization 
studies. Biochim Biphys Acta 1475:125-132. 
27. Perrie Y and Gregoriadis G (2001) Liposome mediated DNA vaccination: The effect of 
vesicle composition. Vaccine 19:3301-3310. 
28. Velinova M, Read N, Kirby C and Gregoriadis G (1996) Morphological observations 
on the fate of liposomes in the regional lymphs nodes after footpad injection into rats. 
Biochim Biophys Acta 1299:207-215. 
29. Szoka FC, Xu Y and Zelpati O (1996) How are nucleic acids released in cells from cationic 
lipid-nucleic acid-complexes? / Lipos Res 6:567-587. 
30. Bacon A, Caparros-Wanderley W, Zadi B and Gregoriadis G (2002) Induction of a cytotoxic 
T lymphocyte (CTL) response to plasmid DNA delivered by Lipodine™. / Lipos 
Res 12:173-183. 
31. Johnson PA, Conwey MA, Daly J, Nicolson C, Robertson J and Mills KH (2000) Plasmid 
DNA encoding influenza virus haemagglutinin induces Th 1 cells and protection against 
respiratory infection despite its limited ability to generate antibody responses. / Gen Virol 
81:1737-1745. 
32. Sha Z, Vincent MJ and Compans RW (1999) (Title) Lmmunobiology 200:21-30. 
33. Gursel M, Tunca S, Ozkan M, Ozcengiz G and Alaeddinoglu G (1999) Immunoadjuvant 
action of plasmid DNA in liposomes. Vaccine 17:1376-1383. 
This page is intentionally left blank
5 
Polymer Micelles as Drug Carriers 
Elena V. Batrakova, Tatiana K. Bronich, 
Joseph A. Vetro and Alexander V. Kabanov 
1. Introduction 
It has long been recognized that improving one or more of the intrinsic adsorption, 
distribution, metabolism, and excretion (ADME) properties of a drug is a critical 
step in developing more effective drug therapies. As early as 1906, Paul Ehrlich 
proposed altering drug distribution by conjugating toxic drugs to "magic bullets" 
(antibodies) having high affinity for cancer cell-specific antigens, in order to both 
improve the therapeutic efficacy of cancer while decreasing its toxicity.1 Since then, 
it has become clear that directly improving intrinsic ADME through modifications 
of the drug is limited or precluded by structural requirements for activity. In other 
words, low molecular mass drugs are too small and have only limited number of 
atomic groups that can be altered to improve ADME, which often adversely affects 
drug pharmacological activity. In turn, the modifications of many low molecular 
mass drugs, aimed to increase their pharmacological activity, often adversely 
affect their ADME properties. For example, the potency and specificity of drugs 
can be improved by the addition of hydrophobic groups.2 The associated decrease 
in water solubility, however, increases the likelihood of drug aggregation, leading 
to poor absorption and bioavailability during oral administration2 or lowered systemic 
bioavailability, high local toxicity, and possible pulmonary embolism during 
intravenous administration.3 
Although there have been considerable difficulties for improving some existing 
drugs through chemical modifications, the problem became even more obvious 
57 
58 Batrakova et al. 
with the development of high-throughput drug discovery technologies. Almost 
half of lead drug candidates identified by high-throughput screening have poor 
solubility in water, and are abandoned before the formulation development stage.4 
In addition, newly synthesized drug candidates often fail due to poor bioavailability, 
metabolism and/or undesirable side effects, which together decrease the 
therapeutic index of the molecules. Furthermore, a new generation of biopharmaceuticals 
and gene therapy agents are emerging based on novel biomacromolecules, 
such as DNA and proteins. The use of these biotechnology-derived drugs is completely 
dependent on efficient delivery to the critical site of the action in the body. 
Therefore, drug delivery research is essential in the translation of newly discovered 
molecules into potent drug candidates and can significantly improve therapies of 
existing drugs. 
Polymer-based drugs and drug delivery systems emerged from the laboratory 
bench in the 1990s as a promising therapeutic strategy for the treatment of certain 
devastating human diseases.5'6 A number of polymer therapeutics are presently 
on the market or undergoing clinical evaluation to treat cancer and other diseases. 
Most of them are low molecular weight drug molecules or therapeutic proteins that 
are chemically linked to water-soluble polymers to increase drug solubility, drug 
stability, or enable targeting to tumors. 
Recently, as a result of rapid development of novel nanotechnology-derived 
materials, a new generation of polymer therapeutics has emerged, using materials 
and devices of nanoscale size for the delivery of drugs, genes, and imaging 
molecules.7-12 These materials include polymer micelles, polymer-DNA complexes 
("polyplexes"), liposomes, and other nanostructured materials for medical use that 
are collectively known as nanomedicines. Compared with first generation polymer 
therapeutics, the new generation nanomedicines are more advanced. They 
entrap small drugs or biopharmaceutical agents such as therapeutic proteins and 
DNA, and can be designed to trigger the release of these agents at the target 
site. Many nanomedicines are constructed using self-assembly principles such as 
the spontaneous formation of micelles or interpolyelectrolyte complexes, driven 
by diverse molecular interactions (hydrophobic, electrostatic, etc.). This chapter 
considers polymeric micelles as an important example of the new generation of 
nanomedicines, which is also perhaps among the most advanced approach toward 
clinical applications in diagnostics and the treatment of human diseases. 
2. Polymer Micelle Structures 
2.1. Self-assembled micelles 
Self-assembled polymer micelles are created from amphiphilic polymers that 
spontaneously form nanosized aggregates when the individual polymer chains 
Polymer Micelles as Drug Carriers 59 
Single polymer chains Polymeric micelle 
("Unimers") 
Fig. 1. Self-assembly of block copolymer micelles. 
("unimers") are directly dissolved in aqueous solution (dissolution method)13 
above a threshold concentration (critical micelle concentration or CMC) and solution 
temperature (critical micelle temperature or CMT) (Fig. 1). Amphiphilic polymers 
with very low water solubility can alternatively be dissolved in a volatile 
organic solvent, then dialyzed against an aqueous buffer (dialysis method).14 
Amphiphilic di-block (hydrophilic-hydrophobic) or tri-block (hydrophilichydrophobic-
hydrophilic) copolymers are most commonly used to prepare selfassembled 
polymer micelles for drug delivery,9'15,16 although the use of graft 
copolymers has been reported.17-19 For drug delivery purposes, the individual 
unimers are designed to be biodegradable20,21 and/or have a low enough molecular 
mass (< ~40 kDa) to be eliminated by renal clearance, in order to avoid polymer 
buildup within the body that can potentially lead to toxicity.22 The most developed 
amphiphilic block copolymers assemble into spherical core-shell micelles approximately 
10 to 80 nm in diameter,23 consisting of a hydrophobic core for drug loading 
and a hydrophilic shell that acts as a physical ("steric") barrier to both micelle 
aggregation in solution, and to protein binding and opsonization during systemic 
administration (Fig. 2). 
The most common hydrophilic block used to form the hydrophilic shell 
is the FDA-approved excipient poly(ethylene glycol) (PEG) or poly(ethylene 
oxide) (PEO).24 PEG or PEO consists of the same repeating monomer subunit 
CH2-CH2-O, and may have different terminal end groups, depending on the 
synthesis procedure, e.g. hydroxyl group HO-(CH2-CH2-0)n-H; methoxy group 
CH30-(CH2-CH2-0)n-H, etc. PEG/PEO blocks typically range from 1 to 15 kDa.16,24 
In addition to its FDA approval, PEG is extremely soluble and has a large 
excluded volume. This makes it especially suitable for physically interfering with 
intra-micelle interactions and subsequent micelle aggregation. PEG also blocks 
protein and cell surface interactions, which greatly decreases nanoparticle uptake 
by the reticuloendothelial system (RES), and consequently increases the plasma 
60 Batrakova etal. 
Self-Assembled 
No self assembly 
Homopolymer 
A n n n 
Di-block copolymer 
Tri-block copolymer 
Graft copolymer 
i n n n***** + ^ ^ , "w>>2^i?^s /^' 
Charged copolymer / ? S 
Covalentlv-Assembled 
(unimolecular micelles) 
Star Dendritic 
hydrophilic block 
hydrophobic block 
cation ic block 
anionic block 
annn 
^ ^ H 
+++++ 
i i i i . 
Fig. 2. Polymer micelle structures. 
half life of the polymer micelle.25 The degree of steric protection by the hydrophilic 
shell is a function of both the density and length of the hydrophilic PEG blocks.25 
Unlike the hydrophilic block, which is typically PEG or PEO, different 
types of hydrophobic blocks have been sufficiently developed as hydrophobic 
drug loading cores.16 Examples of diblock copolymers include (a) poly(L-amino 
acids), (b) biodegradable poly(esters), which includes poly(glycolic acid), poly(D 
lactic acid), poly(D,L-lactic acid), copolymers of lactide/glycolide, and poly(ecaprolactone), 
(c) phospholipids/long chain fatty acids26; and for tri-block 
copolymers, (d) polypropylene oxide (in Pluronics/poloxamers).9 The choice of 
hydrophobic block is largely dictated by drug compatibility with the hydrophobic 
core (when drug is physically loaded, as described later) and the kinetic stability of 
the micelle. 
The self-assembly of amphiphilic copolymers is a thermodynamic and, consequently, 
a reversible process that is entropically driven by the release of ordered 
water from hydrophobic blocks; it is either stabilized or destabilized by solvent 
interactions with the hydrophilic shell. As such, the structural potential of 
amphiphilic copolymer unimers to form micelles is determined by the mass ratio 
of hydrophilic to hydrophobic blocks, which also affects the subsequent morphology 
if aggregates are formed.14 If the mass of the hydrophilic block is too great, 
the copolymers exist in aqueous solution as unimers, whereas, if the mass of the 
hydrophobic block is too great, unimer aggregates with non-micellar morphology 
are formed.27 If the mass of the hydrophilic block is similar or slightly greater than 
the hydrophobic block, then conventional core shell micelles are formed. 
An important consideration for drug delivery is the relative thermodynamic 
(potential for disassembly) and kinetic (rate of disassembly) stability of the polymer 
Polymer Micelles as Drug Carriers 61 
micelle complexes, after intravenous injection and subsequent extreme dilution in 
the vascular compartment.28 This is because the polymer micelles must be stable 
enough to avoid burst release of the drug cargo, as in the case of a physically loaded 
drug, upon systemic administration and remain as nanoparticles long enough to 
accumulate in sufficient concentrations at the target site. 
The relative thermodynamic stability of polymer micelles (which is inversely 
related to the CMC) is primarily controled by the length of the hydrophobic block.13 
An increase in the length of the hydrophobic block alone significantly decreases the 
CMC of the unimer construct (i.e. increases the thermodynamic stability of the polymer 
micelle), whereas an increase in the hydrophilic block alone slightly increases 
the CMC (i.e. decrease the thermodynamic stability of a polymer micelle).14 
Although the CMC indicates the unimer concentration below which polymer 
micelles will begin to disassemble, the kinetic stability determines the rate 
at which polymer micelle disassembly occurs. Many diblock copolymer micelles 
possess good kinetic stability and only slowly dissociate into unimers after extreme 
dilution.29 Thus, although polymer micelles are diluted well below typical unimer 
CMCs29 (10~6-10-7M) after intravenous injection, their relative kinetic stability 
might still be suitable for drug delivery. The kinetic stability depends on several 
factors, including the size of a hydrophobic block, the mass ratio of hydrophilic to 
hydrophobic blocks, and the physical state of the micelle core.14 The incorporation 
of hydrophobic drugs may also further enhance micelle stability. 
2.2. Unimolecular micelles 
Unimolecular micelles are topologically similar to self-assembled micelles, but consist 
of single polymer molecules with covalently linked amphiphile chains. For 
example, copolymers with star-like or dendritic architecture, depending on their 
structure and composition, can either aggregate into multimolecular micelles,30-32 
or exist as unimolecular micelles.33 Dendrimers are widely used as building blocks 
to prepare unimolecular micelles, because they are highly-branched, have welldefined 
globular shape and controled surface functionality.34-40 For example, unimolecular 
micelles were prepared by coupling dendritic hypercores of different 
generations with PEO chains.40'41 The dendritic cores can entrap various drug 
molecules. However, due to the structural limitations involved in the synthesis 
of dendrimers of higher generation, and relatively compact structure of the dendrimers, 
the loading capacity of such micelles is limited. Thus, to increase the loading 
capacity, the dendrimer core can be modified with hydrophobic block, followed 
by the attachment of the PEO chains. For example, Wang et al. recently synthesized 
an amphiphilic 16-arm star polymer with a polyamidoamine dendrimer core 
and arms composed of inner lipophilic poly(e-caprolactone) block and outer PEO 
62 Batrakova et al. 
block.42 These unimolecular micelles were shown to encapsulate a hydrophobic 
drug, etoposide, with high loading capacity. 
Multiarm star-like block copolymers represent another type of unimolecular 
micelles.42-46 Star polymers are generally synthesized by either the arm-first or 
core-first methods. In the arm-first method, monofunctional living linear macromolecules 
are synthesized and then cross-linked either through propagation, using 
a bifunctional comonomer,47 or by adding a multifunctional terminating agent 
to connect precise number of arms to one center.45 Conversely, in the core-first 
method, polymer chains are grown from a multifunctional initiator.43'44'46'48 One 
of the first reported examples of unimolecular micelles, suitable for drug delivery, 
was a three-arm star polymer, composed of mucic acid substituted with fatty acids 
as a lipophilic inner block and PEO as a hydrophilic outer block.44 These polymers 
were directly dispersible in aqueous solutions and formed unimolecular micelles. 
The size and solubilizing capacity of the micelles were varied by changing the ratio 
of the hydrophilic and lipophilic moieties. In addition, star-copolymers with polyelectrolyte 
arms can be prepared to develop pH-sensitive unimolecular micelles as 
drug carriers.46 
2.3. Cross-linked micelles 
The multimolecular micelles structure can be reinforced by the formation of crosslinks 
between the polymer chains. These resulting cross-linked micelles are, in 
essence, single molecules of nanoscale size that are stabile upon dilution, shear 
forces and environmental variations (e.g. changes in pH, ionic strength, solvents 
etc.). There are several reports on the stabilization of the polymer micelles by crosslinking 
either within the core domain49-53 or throughout the shell layer.54-56 In 
these cases, the cross-linked micelles maintained small size and core-shell morphology, 
while their dissociation was permanently suppressed. Stable nanospheres 
from the PEO-b-polylactide micelles were prepared by using polymerizable group 
at the core segment.49 In addition to stabilization, the core polymerized micelles 
readily solubilized rather large molecules such as paclitaxel, and retained high 
loading capacity even upon dilution.50 Formation of interpenetrating network 
of a temperature-sensitive polymer (poly-N-isopropylacrilomide) inside the core 
was also employed for the stabilization of the Pluronic micelles.53 The resulting 
micelle structures were stable against dilution, exhibited temperature-responsive 
swelling behavior, and showed higher drug loading capacity than regular Pluronic 
micelles. 
Recently, a novel type of polymer micelles with cross-linked ionic cores was 
prepared by using block ionomer complexes as templates.57 The nanofabrication of 
these micelles involved condensation of PEO-b-poly(sodium methacrylate) diblock 
Polymer Micelles as Drug Carriers 63 
copolymers by divalent metal cations into spherical micelles of core-shell morphology. 
The core of the micelle was further chemically cross-linked and cations 
removed by dialysis. Resulting micelles represent hydrophilic nanospheres of coreshell 
morphology. The core comprises a network of the cross-linked polyanions 
and can encapsulate oppositely charged therapeutic and diagnostic agents, while a 
hydrophilic PEO shell provides for increased solubility. Furthermore, these micelles 
displayed the pH- and ionic strength-responsive hydrogel-like behavior, due to the 
effect of the cross-linked ionic core. Such behavior is instrumental for the design of 
drug carriers with controled loading and release characteristics. 
3. Drug Loading and Release 
In general, there are three major methods for loading drugs into polymer micelle 
cores: (1) chemical conjugation, (2) physical entrapment or solubilization, and 
(3) polyionic complexation (e.g. ionic binding). 
3.1. Chemical conjuga tion 
Drug incorporation into polymer micelles via chemical conjugation was first proposed 
by Ringsdorf's group58 in 1984. According to this approach, a drug is chemically 
conjugated to the core-forming block of the copolymer via a carefully designed 
pH- or enzyme-sensitive linker, that can be cleaved to release a drug in its active 
form within a cell.59,60 The polymer-drug conjugate then acts as a polymer prodrug 
which self assembles into a core-shell structure. The appropriate choice of 
conjugating bond depends on specific applications. 
The nature of the polymer-drug linkage and the stability of the drug conjugate 
linkage can be controled to influence the rate of drug release, and therefore, the 
effectiveness of the prodrug.61-63 For instance, recent work by Kataoka's group proposed 
pH-sensitive polymer micelles of PEO-b-poly(aspartate hydrazone doxorubicin), 
in which doxorubicin was conjugated to the hydrophobic segments through 
acid-sensitive hydrazone linkers that are stable at extracellular pH 7.4, but degrade 
and release the free drug at acidic pH 5.0 to 6.0 in endosomes and lysosomes.63,64 
The original approach developed by this group used doxorubicin conjugated to the 
poly(aspartic acid) chain of PEO-b-poly(aspartic acid) block copolymer through an 
amide bond.65 Adjusting both the composition of the block copolymer and the concentration 
of the conjugated doxorubicin, led to improved efficacy, as evidenced by 
a complete elimination of solid tumors implanted in mice.66 It was later determined 
that doxorubicin physically encapsulated within the micellar core was responsible 
for antitumor activity. This finding led to the use of PEO-b-poly(aspartate doxorubicin) 
conjugates as nanocontainers for physically entrapped doxorubicin.67 
64 Batrakova et al. 
3.2. Physical entrapment 
The physical incorporation or solublization of drugs within block copolymer 
micelles is generally preferred over micelle-forming polymer-drug conjugates, 
especially for hydrophobic drug molecules. Indeed, many polymers and drug 
molecules do not contain reactive functional groups for chemical conjugation, 
and therefore, specific block copolymers have to be designed for a given type 
of drug. In contrast, a variety of drugs can be physically incorporated into the 
core of the micelles, by engineering the structure of the core-forming segment. In 
addition, molecular characteristics (i.e. molecular weight, composition, presence 
of functional groups for active targeting) within a homologous copolymer series 
can be designed to optimize the performance of a drug for a given drug delivery 
situation.9,14 This concept was introduced by our group in the late 1980s and was 
initially termed "micellar microcontainer",68 but is now widely known as a "micellar 
nanocontainer".9,10 Haloperidol was encapsulated in Pluronic block copolymer 
micelles,68 the micelles were targeted to the brain using brain-specific antibodies 
or insulin, and enhancement of neuroleptic activity by the solubilized drug was 
observed. During the last 25 years, a large variety of amphiphilic block copolymers 
have been explored as nanocontainers for various drugs. 
Different loading methods can be used for physical entrapment of the drug into 
the micelles, including but not limited to dialysis,69-72 oil in water emulsification,69 
direct dissolution,42,73,74 or solvent evaporation techniques.75,76 Depending on the 
method, drug solubilization may occur during or after micelle assembly. The loading 
capacity of the polymer micelles, which is frequently expressed in terms of the 
micelle-water partition coefficient, is influenced by several factors, including both 
the structure of core-forming block and a drug, molecular characteristics of the 
copolymer such as composition, molecular weight, and the solution temperature.13 
Many studies indicate that the most important factor related to the drug solubilization 
capacity of a polymer micelle is the compatibility between the drug and 
the core-forming block.9,14,77-80 For this reason, the choice of the core-forming block 
is most critical. One parameter that can be used to assess the compatibility between 
the polymer and a drug is the Flory-Huggins interaction parameter, Xsp/ defined as 
Xsp= (Ss - <5p)2Vs/kT; where Ss and <5p are Scatchard-Hildebrand solubility parameters, 
and Vs is the molecular volume of the solubilizate. It was successfully used as a 
correlation parameter for the solubilization of aliphatic and aromatic hydrocarbons 
in block copolymer micelles.80,81 Recently, Allen's group82 elegantly demonstrated 
that the calculation and comparison of partial solubility parameters of polymers 
and drugs could be used as a reliable means to predict polymer-drug compatibility 
and to guide formulation development. Polymer micelles, possessing core-forming 
blocks predicted to be compatible with the drug of interest (Ellipticine), were able 
Polymer Micelles as Drug Carriers 65 
to increase the solubility of the drug up to 30,000 times, compared with its saturation 
solubility in water.82 The degree of compatibility between the drug and the 
core-forming block has also been shown to influence the release rate of the drug 
from the micelles. When the environment within the core of the micelle becomes 
more compatible with the drug, it results in a considerable decrease in the rate of 
drug release. 
For a given drug, the extent of incorporation is a function of factors that also 
control the micelle size and/or aggregation number. Such factors include the ratio of 
hydrophobic to hydrophilic block length and the copolymer molecular weight. For 
example, the loading capacity of Pluronic micelles was found to increase with the 
increase in the hydrophobic PPO block length. This effect is attributed to a decrease 
in CMC, and therefore, an increase in aggregation number and micelle core size. 
Also, but to a lesser extent, the hydrophilic block length affects the extent of solubilization, 
such that an increase in percentage of PEO in Pluronic block copolymers 
results in a decrease in the loading capacity of the micelles.80,83-85 For a given ratio of 
PPO-to-PEO, higher molecular weight polymers form larger micelles, and therefore, 
show a higher drug loading capacity. Therefore, the total amount of loaded drug can 
be adjusted as a function of the micellar characteristics as clearly was demonstrated 
by Nagaradjan83 and Kozlov et al.85 Several studies indicate that both the copolymer 
concentration as well as the drug to polymer ratio upon loading, have a complex 
effect on the loading capacity of polymer micelles.79,84,86 In general, more polymer 
chains provide more absorption sites. As a result, solubilization is increased 
with polymer concentration.82 However, the solubilization capacity was found to 
reach a saturation level with an increase of polymer concentration.79 The maximum 
loading level is largely influenced by the interaction between the solubilizate and 
core-forming block, and stronger interactions enable saturation to be reached at 
lower polymer concentration. It was also demonstrated in the studies by Hurter 
and Hatton84'86 that the loading capacity of micelles formed from copolymers with 
high hydrophobic content was independent of the polymer concentration. In addition, 
the location of the incorporated molecules within polymer micelles (micelle 
core or the core-shell interface) determines the extent of solubilization, as well as the 
rate of drug release.87,88 It has been found that more soluble compounds are localized 
at the core-shell interface or even in the inner shell, whereas more hydrophobic 
molecules have a tendency to solubilize in the micelle core.85,87,88 The release rate 
of drug localized in the shell or at the interface appears to account for the "burst 
release" from the micelles.87 In general, for drugs physically incorporated in polymer 
micelles, release is controled by the rate of diffusion of the drug from the micellar 
core, stability of the micelles, and the rate of biodegradation of the copolymer. 
If the micelle is stable and the rate of polymer biodegradation is slow, the diffusion 
rate of the drug will be mainly determined by the abovementioned factors, 
66 Batrakova et al. 
i.e. the compatibility between the drug and core forming block of copolymer,69,82 
the amount of drug loaded, the molecular volume of drug, and the length of the 
core forming block.89 In addition, the physical state of the micelle core and drug 
has a large influence on release characteristics. It was demonstrated that the diffusion 
of incorporated molecules from the block copolymer micelles with glassy 
cores is slower, in comparison to the diffusion out of the cores that are more 
mobile.87 
3.3. Poly ionic complexation 
Charged therapeutic agents can be incorporated into block copolymer micelles, 
through electrostatic interactions with an oppositely charged ionic segment of block 
copolymer. Since it was being proposed independently by Kabanov and Kataoka 
in 1995,90,91 this approach is now widely used for the incorporation of various 
polynucleic acids into block ionomer complexes, for developing non-viral gene 
delivery systems. Ionic block lengths, charge density, and ionic strength of the 
solution affect the formation of stable block ionomer complexes, and therefore, 
control the amount of drug that can be incorporated within the micelles.8'92 The pHand 
salt-sensitivity of such block ionomer micelles provide a unique opportunity to 
control the triggered release of the active therapeutic agent.1563,93-96 Furthermore, 
block ionomer complexes can participate in the polyion interchange reactions which 
are believed to account for the release of the therapeutic agent and DNA in an active 
form inside cells.7 Several comprehensive reviews can be found in the literature that 
focus on block ionomer micelles as drug and gene delivery systems.8,92 In addition, 
physicochemical aspects of the DNA complexes with cationic block copolymers 
have also been recently reviewed.97 
As an example, the metal-complex formation of ionic block copolymer, PEOb-
poly(L-aspartic acid), was explored to prepare polymer micelles incorporating 
cz's-dichlorodiamminoplatinum (II) (CDDP);98,99 a potent chemotherapeutic agent 
widely used in the treatment of a variety of solid tumors, particularly, testicular, 
ovarian, head and neck, and lung tumors.100,101 The CDDP-loaded micelles 
had a size of approximately 20 nm. These micelles showed remarkable stability 
upon dilution in distilled water, while in physiological saline, they displayed sustained 
release of the regenerated Pt complex over 50hrs, due to inverse ligand 
exchange from carboxylate to chloride. The release rate was inversely correlated 
with the chain length of poly(L-aspartic acid) segments in the block copolymer. 
The stability of CDDP-loaded micelle against salt was shown to be improved by 
the addition of homopolymer, poly(L-aspartic acid), in the micelles.102 Recently, 
CDDP-loaded micelles were newly prepared using another block copolymer, 
PEO-b-poly(glutamic acid) to improve and optimize the micellar stability, as well 
Polymer Micelles as Drug Carriers 67 
as the drug release profile.103 The drug loading in the micelles was as high as 39% 
(w/w), and these micelles released the platinum in physiological saline at 37°C in 
sustained manner > 150 hrs, without initial burst of the drug. 
The principle of polyionic complexation can also be used to design new photosensitizers 
for photodynamic therapy of cancer. The group of Kataoka reported 
formation of micelles, as a result of mixing of oppositely charged dendrimer porphyrin 
and block ionomer, based on electrostatic assembly104 or combination of 
electrostatic and hydrogen bonding interactions.95'105 The micelles were stabile at 
physiological conditions and released the entrapped dendrimers in the acidic pH 
environment (pH 5.0), suggesting a possibility of pH-triggered drug release in the 
intracellular endosomal compartments. Overall, the photodynamic efficacy of the 
dendrimer porphyrins was dramatically improved by inclusion into micelles. This 
process resulted in more than two orders of magnitude increase in the photocytotoxicity, 
compared with that of the free dendrimer porphyrins. 
In addition, the polyionic complexation has been used to immobilize charged 
enzymes such as egg white lysozyme106 or trypsin,107 which were incorporated 
in the core of polyion micelles, after mixing with oppositely charged ionic block 
copolymer. A remarkable enhancement of enzymatic activity was observed in 
the core of the micelles. Furthermore, the on-off switching of the enzyme activity 
was achieved through the destabilization of the core domain by applying a 
pulse electric field.108 These unique features of the polyion micelles are relevant 
for their use as smart nanoreactors in the diverse fields of medical and biological 
engineering. 
Last, but not the least, a special class of polyion complexes has been synthesized 
by reacting block ionomers with surfactants of opposite charge, resulting in the 
formation of environmentally responsive nanomaterials, which differ in sizes and 
morphologies, and include micelles and vesicles.109-113 These materials contain a 
hydrophobic core formed by the surfactant tail groups, and a hydrophilic shell 
formed, for example, by PEO chains of the block ionomer. These block ionomer 
complexes can incorporate charged surfactant drugs such as retinoic acid, as well 
as other drugs via solubilization in the hydrophobic domains formed by surfactant 
molecules.114 They display transitions induced by changes in pH, salt concentration, 
chemical nature of low molecular mass counterions, as well as temperature. They 
can also be fine tuned to respond to environmental changes occurring in a very wide 
range of conditions that could realize during delivery of biological and imaging 
agents.94115 The unique self-assembly behavior, the simplicity of the preparation, 
and the wide variety of available surfactant components that can easily produce 
polymer micelles with a very broad range of core properties, make this type of 
materials extremely promising for developing vehicles for the delivery of diagnostic 
and therapeutic modalities. 
68 Batrakova et al. 
4. Pharmacokinetics and Biodistribution 
Incorporation of a low molecular mass drug into polymer micelles drastically alters 
pharmacokinetics and biodistribution of the drug in the body, which is crucial 
for the drug action. Low molecular mass drugs, after administration in the body, 
rapidly extravasate to various tissues affecting them almost indiscriminately, and 
then are rapidly eliminated from the body via renal clearance, often causing toxicity 
to kidneys.116 Furthermore, many drugs display low stability and are degraded in 
the body, often forming toxic metabolites. An example is doxorubicinol, a major 
metabolite of doxorubicin, which causes cardiac toxicity.117 These impediments to 
the therapeutic use of low molecular mass drugs can be mitigated by encapsulating 
drugs in polymer micelles. Within the micelles, the drug molecules are protected 
from enzymatic degradation by the micelle shell. The pharmacokinetics and biodistribution 
of the micelle-incorporated drugs are mainly determined by the surface 
properties, size, and stability of the micelles, and are less affected by the properties 
of the loaded drug. The surface properties of the micelles are determined by 
the micelle shell. The shell from PEO effectively masks drug molecules and prevents 
interactions with serum proteins and cells, which contributes to prolonged 
circulation of the micelles in the body.16 From the size standpoint, polymer micelles 
fit an ideal range of sizes for systemic drug delivery. On the one hand, micelles 
are sufficiently large, usually exceeding 10 nm in diameter, which hinders their 
extravasation in nontarget tissues and prevents renal glomerular excretion. On the 
other hand, the micelles are not considered large, since their size usually does not 
exceed 100 nm. As a result, micelles avoid scavenging by the mononuclear phagocytes 
system (MPS) in the liver and spleen. To this end, "stealth" particles whose 
surface is decorated with PEO are known to be less visible to macrophages and 
have prolonged half-lives in the blood.64,118,119 
The contribution of the micelle stability to pharmacokintetics and biodistribution 
is much less understood, although it is clear that micelle degradation should 
result in a decrease of the size and drug release, perhaps, prematurely. Degradation 
of the micelles, resulting in the formation of block copolymer unimers, could also be 
a principal route for the removal of the polymer material from the body. The molecular 
mass of the unimers of most block copolymers is below the renal excretion limit, 
i.e. less than ~ 20 to 40 kDa,22,120121 while the molecular mass of the micelles, which 
usually contain several dozen or even hundreds of unimers molecules, is above 
this limit. Thus, the unimers are sufficiently small and can be removed via renal 
excretion, while the micelles cannot. A recent study by Batrakova et al. determined 
pharmacokinetic parameters of an amphiphilic block copolymer, Pluronic P85, 
and perhaps provided first evidence that the pharmacokinetic behavior of a block 
copolymer can be a function of its aggregation state.119 Specifically, the formation 
Polymer Micelles as Drug Carriers 69 
of micelles increased the half-life of the block copolymer in plasma and decreased 
the uptake of the block copolymer in the liver. However, it had no effect on the total 
clearance, indicating that the elimination of Pluronic P85 was controled by the renal 
tubular transport of unimers, but not by the rate of micelles disposition or disintegration. 
Furthermore, the values of the clearance suggested that a significant portion 
of the block copolymer was reabsorbed back into the blood, probably, through the 
kidney's tubular membranes. Chemical degradation of the polymers comprising 
the micelles, followed by renal excretion of the relatively low molecular mass products 
of degradation, may be another route for the removal of the micelle polymer 
material from the body. This route could be particularly important in the case of 
the cross-linked or unimolecular micelles, micelles displaying very high stability, 
and / or micelles composed from very hydrophobic polymer molecules that can bind 
and retain considerably biological membranes and other cellular components. 
The delivery of chemotherapeutic drugs to treat tumors is one of the most 
advanced areas of research using polymer micelles. Two approaches have been 
explored to enhance delivery of drug-loaded polymer micelles to the tumor sites: 
(1) passive targeting and (2) vectorized targeting. The passive targeting involves 
enhanced permeability and retention (EPR) effect.122,123 It is based on the fact that 
solid tumors display increased vascular density and permeability caused by angiogenesis, 
impaired lymphatic recovery, and lack of a smooth muscle layer in solid 
tumor vessels. As a result, micellar drugs can penetrate and retain in the sites of 
tumor lesions. At the same time, extravasation of micellar drugs in normal tissues 
is decreased, compared with low molecular drug molecules. Among normal 
organs, spleen and liver can accumulate polymer drugs, but the drugs are eventually 
cleared via the lymphatic system. The increased circulation time of the micellar 
drugs should further enhance exposure of the tumors to the micellar drug, compared 
with the low molecular mass drugs. Along with passive targeting, the delivery 
of micellar drugs to tumors can potentially be enhanced by the modification of 
the surface of the polymer micelles with the targeting molecules, vectors that can 
selectively bind to the surface of the tumor cells. Potential vectors include antibodies, 
aptamers and peptides, capable of binding tumor-specific antigens and other 
molecules diplayed at the surfaces of the tumors.124-126 
Altered biodistribution of a common antineoplastic agent was demonstrated 
for CDDP encapsulated in polyionic micelles with PEO-b-poly(glutamic acid) block 
copolymers.103 Free CDDP is rapidly distributed to each organ, where its levels 
peak at about one hr after i.v. administration. In contrast, in the case of the CDDPincorporated 
micelles, due to their remarkably prolonged blood circulation time, 
the drug level in the liver, spleen and tumor continued to increase up to at least 
24 hrs after injection. Consequently, the CDDP-incorporated micelle exhibited 4-, 
39-, and 20-fold higher accumulation in the liver, spleen and tumor respectively, 
70 Batrakova et al. 
than the free CDDP. At the same time, the encapsulation of CDDP into the micelles 
significantly decreased drug accumulation in the kidney, especially during first hr 
after administration. This suggested potential for the decrease of severe nephrotoxicity 
observed with the free drug, which is excreted through the glomerular 
filtration, thus affecting the kidney.127 
Promising results were also demonstrated for doxorubicin incorporated into 
styrene-maleic acid micelles.128 In this case, as a result of drug entrapment into 
micelles, the drug was redirected from the heart to the tumor, and the doxorubicin 
cardiotoxicity was diminished. Complete blood counts and cardiac histology for 
the micellar drug showed no serious side effects for i.v. doses as high as 100 mg/kg 
doxorubicin equivalent in mice. Similar results were reported for doxorubicin incorporated 
in mixed micelles of PEO-b-poly(L-histidine) and PEO-b-poly(L-lactic acid) 
block copolymers.129 Tissue levels of doxorubicin administered in the micellar formulation 
were decreased in the blood and the liver, and considerably increased in 
the solid tumor, compared with the free drug. Further increase in the tumor delivery 
was achieved by modifying the surface of the micelles with the folate molecules. 
The accumulated doxorubicin levels observed using folate-modified micelles was 
20 times higher than those for free doxorubicin, and 3 times higher than those for 
unmodified micelles. 
The first micellar formulation of doxorubicin to reach clinical evaluation stage, 
used the micelles composed of triblock copolymer, PEO-b-poly (propylene oxide)-b- 
PEO, Pluronic.130 Analysis of pharmacokinetics and biodistribution of doxorubicin 
incorporated into mixed micelles of Pluronics L61 and F127, SP1049C, demonstrated 
more efficient accumulation of the micellar drug in the tumors, compared with the 
free drug. Specifically, the areas under the curves (AUC) in the Lewis lung carcinoma 
3LL M-27 solid tumors in C57B1 / 6 mice were increased about two fold using SP1049, 
compared with the free doxorubicin. Furthermore, this study indicated that the peak 
levels of doxorubicin formulated with SP1049 in the tumor were delayed and the 
drug residence time was increased, in comparison with the free doxorubicin.130 
A clear visualization of drug delivery to the tumor site was shown for doxorubicin 
covalently incorporated through the pH-sensitive link into polymer micelles 
of PEO-poly(aspartate hydrazone doxorubicin).64 A phase-contrast image showed 
that the tumor blood vessels containing the micelles leaked into extra vascular compartments 
of the tumors, resulting in the infiltration of the micelles into tumor 
sites. The micelles circulated in the blood for a prolonged time, and the AUC for 
micellar doxorubicin was 15-fold greater than the AUC for the free doxorubicin. 
Furthermore, the AUC values of the micellar doxorubicin in the heart and kidney 
decreased, compared with the free drug. Thus, the selectivity of drug delivery to 
the tumor, compared with heart and kidney (AUCtumor/AUC0rgan) was increased 
by 6- and 5-folds respectively. This may result in the reduction of side effects of 
Polymer Micelles as Drug Carriers 71 
doxorubicin such as cardiotoxicity and nephrotoxicity. Moreover, the micellar doxorubicin 
showed relatively low uptake in the liver and spleen, despite very long 
residence time in the blood. 
Biodistribution of paclitaxel incorporated into biodegradable polymer micelles 
of monomethoxy-PEO-b-poly(D,L-lactide) block copolymer, Genexol-PM, was 
compared with the regular formulation of the drug in Cremophor EL.131 Two to 
three-fold increases in drug levels were demonstrated in most tissues including 
liver, spleen, kidneys, lungs, heart and tumor, after i.v. administration of Genexol- 
PM, compared with paclitaxel. Nevertheless, acute dose toxicity of Genexol-PM 
was about 25 times lower than that of the conventional drug formulation, which 
appears to be a result of the reformulation avoiding the use of Chremophor EL and 
dehydrated ethanol that are toxic. 
Selective tumor targeting with paclitaxel encapsulated in micelles, modified 
with tumor-specific antibodies 2C5 ("immunomicelles"), was reported using Lewis 
lung carcinoma solid tumor model in C57B1/6J mice.26 These micelles were prepared 
from PEO-distearyl phosphatidylethanolamine conjugates with the free PEO 
end activated with the p-nitrophenylcarbonyl group for the antibody attachment. 
The amount of micellar drug accumulated in the tumor exceeded that in the nontarget 
tissue (muscles) by more than ten times. It is worth noting that the highest 
accumulation in the tumor was demonstrated in the micelles containing the longest 
PEO chains, which also had the longest circulation time in the blood. Furthermore, 
the immunomicelles displayed the highest amount of tumor-accumulated drug, 
compared with either free paclitaxel or non-vectorized micelles. It was demonstrated 
that paclitaxel delivered by plain micelles in the interstitial space of the 
tumor was eventually cleared after gradual micellar degradation. In contrast, 
paclitaxel-loaded 2C5 immunomicelles were internalized by cancer cells and the 
retention of the drug inside the tumor was enhanced.132 
Unexpected results were found using pH-sensitive polymer micelles of Nisopropylacrylamide 
and methacrylic acid copolymers randomly or terminally 
alkylated with octadecyl groups.64,133 It was demonstrated that aluminium chloride 
phthalocyanine (AlClPc) incorporated in such micelles was cleared more rapidly 
and less accumulated in the tumor, than the AlClPc formulated with Cremophor 
EL. Furthermore, significant accumulation in the liver and spleen (and lungs for 
most hydrophobic copolymers) was observed, compared with Cremophor EL formulation. 
The enhanced uptake of such polymer micelles by the cells of mononuclear 
phagocyte system (MPS) could be due to micelle aggregation in the blood 
and embolism in the capillaries. Thus, it attempted to reduce the uptake of the 
micelles in MPS by incorporating water soluble monomers, N-vinyl-2-pyrrolidone 
in the copolymer structure.134 The modified formulation displayed same levels of 
tumor accumulation and somewhat higher antitumor activity than the Cremophor 
72 Batrakova et al. 
EL formulation. This work serves as an example reinforcing the need of proper 
adjustment of the polymer micelle structure, and perhaps the need of using block 
copolymers to produce a defined protective hydrophilic shell to facilitate evasion 
of the polymer micelles from MPS. 
5. Drug Delivery Applications 
The studies on the application of polymer micelles in drug delivery have mostly 
focused on the following areas that are considered below: (1) delivery of anticancer 
agents to treat tumors; (2) drug delivery to the brain to treat neurodegenerative diseases; 
(3) delivery of antifungal agents; (4) delivery of imaging agents for diagnostic 
applications; and (5) delivery of polynucleotide therapeutics. 
5.1. Chemotherapy of cancer 
To enhance chemotherapy of tumors using polymer micelles, four major approaches 
were employed: (1) passive targeting of polymer micelles to tumors due to EPR 
effect; (2) targeting of polymer micelles to specific antigens overexpressed at the 
surface of tumor cells; (3) enhanced drug release at the tumor sites having low pH; 
and (4) sensitization of drug resistant tumors by block copolymers. 
A series of pioneering studies by Kataoka's group used polymer micelles for 
passive targeting of various anticancer agents and chemotherapy of tumors.102,103'135 
One notable recent example reported by this group involves polymer micelles of 
PEO-b-poly(L-aspartic acid) incorporating CDDP. Evaluation of anticancer activity 
using murine colon adenocarcinoma C26 as an in vivo tumor model, demonstrated 
that CDDP in polymer micelles had significantly higher activity than the free CDDP, 
resulting in complete eradication of the tumor.103 A formulation of paclitaxel in 
biodegradable polymer micelles of monomethoxy-PEO-b-poly(D,L-lactide) block 
copolymer, Genexol-PM, also displayed elevated activity in vivo against human 
ovarian carcinoma OVCAR-3 and human breast carcinoma MCF7, compared with 
a regular formulation of the drug in Cremophor EL.131 In addition, anthracycline 
antibiotics, doxorubicin and pirarubicin, incorporated in styrene-maleic acid 
micelles each revealed potent anticancer effects in vivo against mouse sarcoma 
S-180, resulting in complete eradication of tumors in 100% of tested animals.128 
Notably, animals survived for more than one year, after treatment with the micelleincorporated 
pirarubicin at doses as high as lOOmg/kg of pirarubicin equivalent. 
Complete blood counts, liver function test, and cardiac histology showed no 
sign of adverse effects for intravenous doses of the micellar formulation. In contrast, 
animals receiving free pirarubicin had a much reduced survival and showed 
serious side effects.136 Collectively, these studies suggested that various micelleincorporated 
drugs display improved therapeutic index in solid tumors, which 
Polymer Micelles as Drug Carriers 73 
correlates with enhanced passive targeting of the drug to the tumor sites, as well as 
decreased side effects, compared with conventional formulations of these drugs. 
Tumor-specific targeting of polymer micelles to molecular markers expressed 
at the surface of the cancer cells has also been explored to eradicate tumor cells. 
For example, a recent study by Gao's group developed a polymer micelle carrier 
to deliver doxorubicin to the tumor endothelial cells with overexpressed Xvfi3 
integrins.137 A cyclic pentapeptide, cRGD was used as a targeting ligand that is 
capable of selective and high affinity binding to the Xvfio, integrin. Micelles of PEOb-
poly(e-caprolactone) loaded with doxorubicin were covalently bound with cRGD. 
As a result of such modification, the uptake of doxorubicin-containing micelles in 
in vitro human endothelial cell model derived from Kaposi's sarcoma, was profoundly 
increased. In addition, folate receptor often overexpressed in cancer cells 
has been evaluated for targeting various drug carriers to tumors.138 This strategy 
has also been evaluated to target polymer micelles. For example, mixed micelles 
of PEO-b-poly(L-histidine) and PEO-b-poly(L-lactic acid) block copolymers with 
solubilized doxorubicin129 or micelles of PEO-b-poly(DL-lactic-co-glycolic acid) 
block copolymer with covalently attached doxorubicin,139 were each surface modified 
by conjugating folate molecules to the free PEO ends. In both cases, in vitro 
and in vivo studies demonstrated increased antitumor activity of the micelleincorporated 
drug resulting from such modification. The enhanced delivery of the 
micellar drugs through the folate receptor, and the enhanced retention of the modified 
micelles at the tumor sites are possible explanations for the effects of these folate 
modifications. 
Micelles conjugated with antibodies or antibody fragments capable to 
recognize tumor antigens were shown to improve therapeutic efficacy in vivo over 
non-modified micelles.23 This approach can result in high selectivity of binding, 
internalization, and effective retention of the micelles in the tumor cells. In addition, 
recent advances in antibody engineering allow for the production of humanized 
antibody fragments, reducing problems with immune response against mouse 
antibodies.140 For example, micelles of PEO-distearyl phosphatidylethanolamine 
were covalently modified with the monoclonal antibody 2C5 that binds to microsomes, 
displayed at the surface of many tumor cells. The micelles were then 
used for incorporating various poorly soluble anticancer drugs including tamoxifen, 
paclitaxel, dequalinium, and chlorine e6 trimethyl ester.26'132'141 It was shown 
that paclitaxel-loaded 2C5-immunomicelles could specifically recognize a variety 
of tumor types. The binding of these immunomicelles was observed for all 
cancer cell lines tested, i.e. murine Lewis lung carcinoma, T-lymphoma EL4, and 
human breast adenocarcinomas, BT-20 and MCF7.141 Moreover, paclitaxel-loaded 
2C5 immunomicelles demonstrated highest anticancer activity in Lewis lung carcinoma 
tumor model in mice, compared with plain paclitaxel-loaded micelles and 
74 Batrakova et al. 
the free drug.132 The increased antitumor effect of immunomicelles in vivo correlated 
with the enhanced retention of the drug delivered with the immunomicelles 
inside the tumor. 
Tumors often display low pH of interstitial fluid, which is mainly attributed 
to higher rates of aerobic and anaerobic glycolysis in cancer cells than in normal 
cells.142,143 This phenomenon has been employed in the design of various 
pH-sensitive polymer micelle systems for the delivery of anticancer drugs to the 
tumors. One approach consisted in the chemical conjugation of anticancer drugs 
to the block copolymers through pH sensitive cleavable links that are stable at 
neutral pH, but are cleavable and release the drug in the mildly acidic pH. For 
example, several groups used hydrasone-based linking groups, to covalently attach 
doxorubicin to PEO-b-poly(DL-lactic-co-glycolic acid) block copolymer,21,144 PEOb-
block-poly(allyl glycidyl ether)145 or PEO-b-poly(aspartate hydrazone) block 
copolymer.63,64 It was suggested that doxorubicin will remain in the micelles in 
the blood stream, and will be released at tumor sites at lower pH. For example, 
in vitro and in vivo studies using PEO-b-poly(aspartate hydrazone doxorubicin) 
micelles demonstrated that the micelles display an intracellular pH-triggered drug 
release capability, tumor-infiltrating permeability, and effective antitumor activity 
with extremely low toxicity.63,64 Overall, the animal studies suggested that such 
polymer micelle drug has a wide therapeutic window due to increased efficacy 
and decreased toxicity, compared with free doxorubicin.64 
An alternative mechanism for pH-induced triggering of drug release at the 
tumor sites consists of using pH sensitive polyacids or polybases as building 
blocks for polymer micelles.94,146,147 For example, mixed micelles of PEO-bpoly(
L-histidine) and PEO-b-poly(L-lactic acid) block copolymers incorporate pHsensitive 
poly-base, poly(L-histidine) in the hydrophobic core.147 The core can also 
solubilize hydrophobic drugs such as doxorubicin. The protonation of the polybase 
at acidic conditions resulted in the destabilization of the core and triggered 
release of the drug. This system was also targeted to the tumors through the folate 
molecules as described earlier and has shown significant in vivo antitumor activity 
and less side effects, compared with the free drug.129 Notably, it was also effective 
in vitro and in vivo against multidrug resistant (MDR) human breast carcinoma 
MCF7/ADR that overexpresses P-glycoprotein (Pgp). Pgp is a drug efflux transport 
protein that serves to eliminate drugs from the cancer cells and significantly 
decreases the anticancer activity of the drugs. The micelle incorporated drug was 
released inside the cells, and thus avoided the contact with Pgp localized at the 
cell plasma membrane, which perhaps contributed to the increased activity of pH 
sensitive doxorubicin micelles in the MDR cells. 
A different approach using Pluronic block copolymer micelles to overcome 
MDR in tumors has been developed by our group.130,148-151 Studies by Alakhov 
Polymer Micelles as Drug Carriers 75 
et al. demonstrated that Pluronic block copolymers can sensitize MDR cells, 
resulting in an increased cytotoxic activity of doxorubicin, paclitaxel, and other 
drugs by 2,3 orders of magnitude.148'149 Remarkably, Pluronic can enhance drug 
effects in MDR cells through multiple effects including (1) inhibiting drug efflux 
transporters, such as Pgp149-152 and multidrug resistance proteins (MRPs),153'154 
(2) abolishing drug sequestration within cytoplasmic vesicles,149'153 (3) inhibiting 
the glutathione/glutathione S-transferase detoxification system,154 and (4) enhancing 
proapoptotic signaling in MDR cells.155 Similar effects of Pluronics have also 
been reported using in vivo tumor models.130,150 In these studies, mice bearing 
drug-sensitive and drug-resistant tumors were treated with doxorubicin alone 
and with doxorubicin in Pluronic compositions. The tumor panel included i.p. 
murine leukemias (P388, P388-Dox), s.c. murine myelomas (Sp2/0, Sp2/0-Dnr), 
i.v. and s.c. Lewis lung carcinoma (3LL-M27), s.c. human breast carcinomas (MCF7, 
MCF7/ADR), and s.c. human oral epidermoid carcinoma (KBv).130 Using the NCI 
criteria for tumor inhibition and increased lifespan, Pluronic/doxorubicin has met 
the efficiency criteria in all models (9 of 9), while doxorubicin alone was only effective 
in selected tumors (2 of 9) .130 Results showed that the tumors were more responsive 
in the Pluronic /doxorubicin treatment groups than in doxorubicin alone. These 
studies demonstrated improved treatment of drug resistant cancers with Pluronics. 
The mechanisms of effects of Pluronic on Pgp have been studied in great 
detail.151 In particular, exposure of MDR cells to Pluronics has resulted in the 
inhibition of Pgp-mediated efflux,149 and this overcomes defects in intracellular 
accumulation of Pgp-dependent drugs,148,149,152 and abolishes the directionality 
difference in the flux of these drugs across polarized cell monolayers.156-158 
The lack of changes in membrane permeability with Pluronics to (1) non-Pgp 
compounds in MDR cells,158,159 and (2) Pgp probes in non-MDR cells,149,153 suggested 
that Pluronic effects were specific to the Pgp efflux system. These effects 
were observed at Pluronic concentrations less than or equal to the critical micelle 
concentration (CMC).152,159 Thus, Pluronic unimers rather than the micelles were 
responsible for these effects. Specifically, Pluronic molecules displayed a dual function 
in MDR cells.160-162 Firstly, they incorporated into the cell membranes and 
decreased the membrane microviscosity. This was accompanied by the inhibition 
of Pgp ATPase activity. Secondly, they translocated into cells and reached intracellular 
compartments. This was accompanied by the inhibition of respiration,163 
presumably due to Pluronic interactions with the mitochondria membranes. As a 
result, within 15 min after exposure to select Pluronics, intracellular levels of ATP in 
MDR cells were drastically decreased.160-162 Remarkably, such ATP depletion was 
not observed in non-MDR cells, suggesting that the Pluronic was "selective", with 
respect to the MDR phenotype.160'164 Combining these two effects, Pgp ATPase inhibition 
and ATP depletion, resulted in the shut-down of the efflux system in MDR 
76 Batrakova et al. 
cells.160-162 The Pgp remained functionally active when (1) ATP was restored using 
an ATP supplementation system in the presence of a Pluronic, or (2) when ATP was 
depleted, but there was no direct contact between the Pluronic and Pgp (and no 
ATPase inhibition). Overall, these detailed studies which resulted in the development 
of a micellar formulation of doxorubicin that is evaluated clinically, reinforce 
the fact that block copolymers, comprising the micelles, can serve as biological 
response modifying agents that can have beneficial effects in the chemotherapy of 
tumors. 
5.2. Drug delivery to the brain 
By restricting drug transport to the brain, the blood brain barrier (BBB) represents a 
formidable impediment for the treatment of brain tumors and neurodegenerative 
diseases such as HIV-associated dementia, stroke, Parkinson's and Alzheimer's 
diseases. Two strategies using polymer micelles have been evaluated to enhance 
delivery of biologically active agents to the brain. The first strategy is based on 
the modification of polymer micelles with antibodies or ligand molecules capable 
of transcytosis across brain microvessel endothelial cells, comprising the BBB. The 
second strategy uses Pluronic block copolymers to inhibit drug efflux systems, 
particularly, Pgp, and selectively increase the permeability of BBB to Pgp substrates. 
An earlier study used micelles of Pluronic block copolymers for the delivery 
of the CNS drugs to the brain.68'73 These micelles were surface-modified by attaching 
to the free PEO ends, either polyclonal antibodies against brain-specific antigen, 
a2-glycoprotein, or insulin to target the receptor at the lumenal side of BBB. 
The modified micelles were used to solubilize fluorescent dye or neuroleptic drug, 
haloperidol, and these formulations were administered intravenously in mice. Both 
the antibody and insulin modification of the micelles resulted in enhanced delivery 
of the fluorescent dye to the brain and drastic increases in neuroleptic effect of 
haloperidol in the animals. Subsequent studies using in vitro BBB models demonstrated 
that the micelles, vectorized by insulin, undergo receptor-mediated transport 
across brain microvessel endothelial cells.156 Based on one of these observations, 
one should expect development of novel polymer micelles that target specific 
receptors at the surface of the BBB to enhance transport of the incorporated drugs 
to the brain. 
The studies by our group have also demonstrated that selected Pluronic block 
copolymers, such as Pluronic P85, are potent inhibitors of Pgp, and they have the 
increased entry of the Pgp-substrates to the brain across BBB.156'158'159'165 Pluronic 
did not induce toxic effect in BBB, as revealed by the lack of alteration in paracellular 
permeability of the barrier,156'158 and in histological studies, using specific markers 
for brain endothelial cells.166 Overall, this strategy has potential in developing 
Polymer Micelles as Drug Carriers 77 
novel modalities for the delivery of various drugs to the brain, including selective 
anti cancer agents to treat metastatic brain tumors, as well as HIV protease 
inhibitors to eradicate HIV virus in the brain.167'168 
5.3. Formulations of antifungal agents 
The need for safe and effective modalities for the delivery of chemotherapeutic 
agents to treat systemic fungal infections in immunocompromised AIDS, surgery, 
transplant and cancer patients is very high. The challenges to the delivery of antifungal 
agents include low solubility and sometimes high toxicity of these agents. 
These agents, such as amphotericin B, have low compatibility with hydrophobic 
cores of polymer micelles formed by many conventional block copolymers. Thus, to 
increase solubilization of amphotericin B, the core-forming blocks of methoxy-PEOb-
poly(L-aspartate) were derivatized with stearate side chains.169-172 The resulting 
block copolymers formed micelles. Amphotericin B interacted strongly with the 
stearate side chains in the core of the micelles, resulting in an efficient entrapment 
of the drug in the micelles, as well as subsequent sustained release in the external 
environment. As a result of solubilization of amphotericin B in the micelles, the 
onset of hemolytic activity of this drug toward bovine erythrocytes was delayed, 
relative to that of the free drug.171 Using a neutropenic murine model of disseminated 
Candidas, it was shown that micelle-incorporated amphotericin B retained 
potent in vivo activity. Pluronic block copolymers were used by the same group 
for incapsulation of another poorly soluble antifungal agent, nystatin.172 This is a 
commercially available drug that has shown potential for systemic administration, 
but has never been approved for that purpose, due to toxicity issues. The possibility 
to use Pluronic block copolymers to overcome resistance to certain antifungal 
agents has also been demonstrated.173-176 Overall, one should expect further scientific 
developments using polymer micelle delivery systems for the treatment of 
fungal infection. 
5.4. Delivery of imaging agents 
Efficient delivery of imaging agents to the site of disease in the body can improve 
early diagnostics of cancer and other diseases. The studies in this area using polymer 
micelles as carriers for imaging agents were initiated by Torchilin.177 For example, 
micelles of amphiphilic PEO-lipid conjugates were loaded with i n In and 
gadolinium diethylenetriamine pentaacetic acid-phosphatidylethanolamine (Gd- 
DTPA-PE) and then used for visualization of local lymphatic chain after subcutaneous 
injection into the rabbit's paw.178 The images of local lymphatics were 
acquired using a gamma camera and a magnetic resonance (MR) imager. The 
78 Batrakova et al. 
injected micelles stayed within the lymph fluid, thus serving as lymphangiographic 
agents for indirect MR or gamma lymphography. Another polymer micelle system 
composed of amphiphilic methoxy-PEO-b-poly[epsilon,N-(triiodobenzoyl)-Llysine] 
block copolymers, labeled with iodine, was administered systemically in 
rabbits and visualized by X-ray computed tomography.179 The labeled micelles 
displayed exceptional 24 hrs half-life in the blood, which is likely due to the coreshell 
architecture of the micelle carriers that protected the iodine-containing core. 
Notably, small polymer micelles (<20nm) may be advantageous for bioimaging 
of tumors, compared with PEG-modified long-circulating liposomes (ca. lOOnm). 
In particular, the micelles from PEO-distearoyl phosphatidyl ethanolamine conjugates 
containing m In-labeled model protein were more efficacious in the delivery of 
protein to Lewis lung carcinoma than larger long-circulating liposomes.180 Overall, 
polymer micelles loaded with various agents for gamma, magnetic resonance, and 
computed tomography imaging represent promising modalities for non-invasive 
diagnostics of various diseases. 
5.5. Delivery of polynucleotides 
To improve the stability of polycation-based DNA, delivery complexes in dispersion 
block and graft copolymers containing segments from polycations and nonionic 
water-soluble polymers, such as PEO, were developed.90,181,182 Binding of 
these copolymers with DNA results in the formation of micelle-like block ionomer 
complexes ("polyion complex micelles"), containing hydrophobic sites formed by 
the polycation-neutralized DNA and hydrophilic sites formed by the PEO chains. 
Despite neutralization of charge, complexes remain stable in aqueous dispersion 
due to the effect of the PEO chains.183 Overall, the PEO modified polycation-DNA 
complexes form stable dispersions and do not interact with serum proteins.183,184 
These systems were successfully used for intravitreal delivery of an antisense 
oligonucleotide and the suppression of gene expression in retina in rats.185 Furthermore, 
they displayed extended plasma clearance kinetics and were shown to 
transfect liver and tumor cells, after systemic administration in the body.186-188 In 
addition, there is a possibility targeting such polyplexes to the specific receptors at 
the surface of the cell, for example, by modifying the free ends of PEO chains with 
specific targeting ligands.189-191 Alternatively, to increase the binding of the complexes 
with the cell membrane and the transport of the polynucleotides inside the 
cells, the polycations were modified with amphiphilic Pluronic molecules.192,193 One 
recent study has shown a potential of Pluronic-polyethyleneimine-based micelles 
for in vivo delivery of antisense oligonucleotides to tumors, and have demonstrated 
sensitization of the tumors to radiotherapy as a result of systemic administration 
of the oligonucleotide-loaded micelles.194 
Polymer Micelles as Drug Carriers 79 
6. Clinical Trials 
Three polymer micelle formulations of anticancer drugs have been reported to 
reach clinical trials. The doxorubicin-conjugated polymer micelles developed by 
Kataoka's group195 have progressed recently to Phase I clinical trial at the National 
Cancer Center Hospital, Tokyo, Japan. The micelle carrier NK911 is based on PEO-bpoly(
aspartic acid) block copolymers, in which the aspartic acid units were partially 
(ca. 45%) substituted with doxorubicin to form hydrophobic block. The resulting 
substituted block copolymer forms micelles that are further noncovalently loaded 
with free doxorubicin. Preclinical studies in mice demonstrated higher NK911 activity 
against Colon 26, M5076, and P388, compared with the free drug. Moreover, 
NK911 has less side effects, resulting in less animal body and toxic death than the 
free drug.196 
Clinically, the Pluronic micelle formulation of doxorubicin has been most 
advanced. Based on the in vivo efficacy evaluation, Pluronic L61 was selected for 
clinical development for the treatment of MDR cancers. The final block copolymer 
formulation is a mixture of 0.25% Pluronic L61 and 2% Pluronic F127, formulated 
in isotonic buffered saline.130 This system contains mixed micelles of L61 and F127, 
with an effective diameter of ca. 22 to 27 nm and is stable in the serum. Prior to 
administration, doxorubicin is mixed with this system, which results in spontaneous 
incorporation of the drug in the micelles. The drug is easily released by diffusion 
after dilution of the micelles. The formulation of doxorubicin with Pluronic, 
SP1049C, is safe, following systemic administration based on toxicity studies in 
animals.130 A two-site Phase I clinical trial of SP1049C has been completed.197 Based 
on its results, the dose-limiting toxicity of SP1049C was myelosuppression, reached 
at 90mg/m2 (maximum tolerated dose was 70mg/m2). Phase II study of this formulation 
to treat inoperable metastatic adenocarcinoma of the esophagus is near 
completion as well.198 
Finally, Phase I studies were reported for Genexol-PM, a Cremophor-free polymer 
micelle-formulated paclitaxel.199 Twenty-one patient entered into this study 
with lung, colorectal, breast, ovary, and esophagus cancers. No hypersensitivity 
reaction was observed in any patient. Neuropathy and myalgia were the most 
common toxicities. There were 14% partial responses. The paclitaxel area under 
the curve and peak of the drug concentration in the blood were increased with the 
escalating dose, suggesting linear pharmacokinetics for Genexol-PM.199 
7. Conclusions 
Approximately two decades have passed since the conception of the polymer 
micelle conjugates and nanocontainers for drug delivery. During the first decade, 
80 Batrakova et al. 
only a few studies were published; however, more recently, the number of publications 
in this field has increased tremendously. During this period, novel biocompatible 
and/or biodegradable block copolymer chemistries have been researched, 
the block ionomer complexes capable of incorporating DNA and other charged 
molecules have been discovered, the pH and other chemical signal sensitive micelles 
have been developed. Many studies focused on the use of polymer micelles for 
delivery of poorly soluble and toxic chemotherapeutic agents to the tumors to 
treat cancer. There has been considerable advancement in understanding the processes 
of polymer micelle delivery into the tumors, including passive and vectorized 
targeting of the polymer micelles. Notable achievements also include the studies 
demonstrating the possibilities for overcoming multidrug resistance in cancer, and 
enhancing drug delivery to the brain using block copolymer micelles systems. Overall, 
it is clear that this area has reached a mature stage, reinforced by the fact that 
several human clinical trials using polymer micelles for cancer drug delivery have 
been initiated. At the same time, it is obvious that the possibilities for delivery of 
the diagnostic and therapeutic agents using polymer micelles are extremely broad, 
and one should expect further increase in the laboratory and clinical research in 
this field during the next decade. Targeting polymer micelles to cancer sites within 
the body will address an urgent need to greatly improve the early diagnosis and 
treatment of cancer. Capabilities for the discovery and use of targeting molecules 
will support the development of multifunctional therapeutics that can carry and 
retain antineoplastic agents within tumors. This will also be instrumental in developing 
novel biosensing and imaging modalities for the early detection of cancer 
and other devastating human diseases. 
Acknowledgment 
The authors acknowledge the support of the research using polymer micelles by 
grants from the National Institutes of Health CA89225, NS36229 and EB000551, 
as well as the National Science Foundation DMR0071682, DMR0513699 and BES- 
9907281. We also acknowledge financial support of Supratek Pharma, Inc. (Montreal, 
Canada). AVK and EVB are shareholders and AVK serves as a consultant to 
this Company. 
References 
1. Ehrlich P (1956) The relationship existing between chemical constitution, distribution, 
and pharmacological action, in Himmelweite F, Marquardt M and Dale H (eds.) The 
Collected Papers of Paul Ehrlich. Pergamon, Elmsford, New York, Vol. 1, pp. 596-618. 
2. Lipinski CA, Lombardo F, Dominy BW and Feeney PJ (2001) Experimental and computational 
approaches to estimate solubility and permeability in drug discovery and 
development settings. Adv Drug Del Rev 46:3-26. 
Polymer Micelles as Drug Carriers 81 
3. Fernandez AM, Van Derpoorten K, Dasnois L, Lebtahi K, Dubois V, Lobl TJ, Gangwar 
S, Oliyai C, Lewis ER, Shochat D and Trouet A (2001) N-Succinyl-(beta-alanyl-L-leucyl- 
L-alanyl-L-leucyl)doxorubicin: An extracellularly tumor-activated prodrug devoid of 
intravenous acute toxicity. / Med Chem 44:3750-3753. 
4. Thompson TN (2001) Optimization of metabolic stability as a goal of modern drug 
design. Med Res Rev 21:412-449. 
5. Langer R (2001) Drag delivery. Drags on target. Science 293:58-59. 
6. Duncan R (2003) The dawning era of polymer therapeutics. Nat Rev Drug Discov 
2:347-360. 
7. Kabanov AV and Kabanov VA (1995) DNA complexes with polycations for the delivery 
of genetic material into cells. Bioconjug Chem 6:7-20. 
8. Kakizawa Y and Kataoka K (2002) Block copolymer micelles for delivery of gene and 
related compounds. Adv Drug Del Rev 54:203-222. 
9. Kabanov AV and Alakhov VY (2002) Pluronic block copolymers in drug delivery: From 
micellar nanocontainers to biological response modifiers. Crit Rev Ther Drug Can Syst 
19:1-72. 
10. Savic R, Luo L, Eisenberg A and Maysinger D (2003) Micellar nanocontainers distribute 
to defined cytoplasmic organelles. Science 300:615-618. 
11. Hubbell JA (2003) Materials science. Enhancing drug function. Science 300:595-596. 
12. Salem AK, Searson PC and Leong KW (2003) Multifunctional nanorods for gene delivery. 
Nat Mater 2:668-671. 
13. Kabanov AV, Nazarova IR, Astafieva IV, Batrakova EV, Alakhov VY, Yaroslavov AA 
and Kabanov VA (1995) Micelle formation and solubilization of fluorescent probes 
in poly(oxyethylene-b-oxypropylene-b-oxyethylene) solutions. Macromolecules 28: 
2303-2314. 
14. Allen C, Maysinger D and Eisenberg A (1999) Nano-engineering block copolymer 
aggregates for drug delivery. Coll Surf B: Biointerf 16:3-27. 
15. Kataoka K, Harada A and Nagasaki Y (2001) Block copolymer micelles for drug delivery: 
Design, characterization and biological significance. Adv Drug Del Rev 47:113-131. 
16. Adams ML, Lavasanifar A and Kwon GS (2003) Amphiphilic block copolymers for 
drug delivery. J Pharm Sci 92:1343-1355. 
17. Sakuma S, Hayashi M and Akashi M (2001) Design of nanoparticles composed of graft 
copolymers for oral peptide delivery. Adv Drug Del Rev 47:21-37. 
18. Francis MF, Lavoie L, Winnik FM and Leroux JC (2003) Solubilization of cyclosporin A 
in dextran-g-polyethyleneglycolalkyl ether polymeric micelles. Eur } Pharm Biopharm 
56:337-346. 
19. Francis MF, Piredda M and Winnik FM (2003) Solubilization of poorly water soluble 
drugs in micelles of hydrophobically modified hydroxypropylcellulose copolymers. 
/ Control Rel 93:59-68. 
20. Jeong B, Bae YH and Kim SW (2000) Drug release from biodegradable injectable thermosensitive 
hydrogel of PEG-PLGA-PEG triblock copolymers. / Control Rel 63:155-163. 
21. Yoo HS and Park TG (2001) Biodegradable polymeric micelles composed of doxorubicin 
conjugated PLGA-PEG block copolymer. / Control Rel 70:63-70. 
82 Batrakova et al. 
22. Duncan R and Kopecek J (1984) Soluble synthetic polymers as potential drug carriers. 
AdvPolym Sci 57:51-101. 
23. Torchilin VP (2004) Targeted polymeric micelles for delivery of poorly soluble drugs. 
Cell Mol Life Sci 61:2549-2559. 
24. Torchilin VP (2001) Structure and design of polymeric surfactant-based drug delivery 
systems. / Control Rel 73:137-172. 
25. Moghimi SM, Hunter AC and Murray JC (2001) Long-circulating and target-specific 
nanoparticles: Theory to practice. Pharmacol Rev 53:283-318. 
26. Lukyanov AN, Gao Z and Torchilin VP (2003) Micelles from polyethylene glycol/
phosphatidylethanolamine conjugates for tumor drug delivery. / Control Rel 
91:97-102. 
27. Zhang L, Yu K and Eisenberg A (1996) Ion-Induced morphological changes in "crewcut" 
aggregates of amphiphilic block copolymers. Science 272:1777-1779. 
28. Kabanov AV and Alakhov VY (2000) Micelles of amphilphilic block copolymers as 
vehicles for drug delivery, in, Alexandridis P and Lindman B (eds.) Amphiphilic Block 
Copolymers: Self-Assembly and Applications: Elsevier: Amsterdam, Lausanne, New York, 
Oxford, Shannon, Singapore, Tokyo, pp. 347-376. 
29. Kwon GS and Okano T (1999) Soluble self-assembled block copolymers for drug delivery. 
Pharm Res 16:597-600. 
30. Kim KH, Guo HC, Lim HJ, Huh J, Ahn C-H and Jo WH (2004) Synthesis and micellization 
of star-shaped poly (ethylene glycol)-block-poly(e-caprolactone). Macromolec Chem 
Phys 205:1684-1692. 
31. Gitsov I and Frechet JMJ (1993) Solution and solid-state properties of hybrid lineardendritic 
block copolymers. Macromolecules 26:6536-6546. 
32. Gitsov I, Lambrych KR, Remnant VA and Pracitto R (2000) Micelles with highly 
branched nanoporous interior: Solution properties and binding capabilities of 
amphiphilic copolymers with linear dendritic architecture. / Polym Sci Part A: Polym 
Chem 38:2711-2727. 
33. Hawker CJ, Wooley KL and Frechet JMJ (1993) Unimolecular micelles and globular 
amphiphiles: Dendritic macromolecules as novel recyclable solubilization agents. 
/ Chem Soc Perkin Trans 1: Org Bio-Org Chem (1972-1999):1287-1297. 
34. Tomalia DA, Berry V, Hall M and Hedstrand DM (1987) Starburst dendrimers. 4. 
Covalently fixed unimolecular assemblages reiminiscent of spheroidal micelles. Macromolecules 
20:1164-1167. 
35. Stevelmans S, Hest JCMv, Jansen JFGA, Van Boxtel DAFJ, de Berg EMM and Meijer EW 
(1996) Synthesis, characterization, and guest-host properties of inverted unimolecular 
dendritic micelles. J Am Chem Soc 118:7398-7399. 
36. van Hest JCM, Delnoye DAP, Baars MWPL, van Genderen MHP and Meijer EW (1995) 
Polystyrene-dendrimer amphiphilic block copolymers with a generation-dependent 
aggregation. Science 268:1592-1595. 
37. van Hest JCM, Elissen-Roman C, Baars MWPL, Delnoye DAP, Van Genderen MHP and 
Meijer EW (1995) Polystyrene-poly(propylene imine) dendrimer block copolymers: A 
new class of amphiphiles. Polym Mater Sci Eng 73:281-282. 
Polymer Micelles as Drug Carriers 83 
38. Lorenz K, Muelhaupt R, Frey H, Rapp U and Mayer-Posner FJ (1995) Carbosilane-Based 
Dendritic Polyols. Macromolecules 28:6657-6661. 
39. Gitsov I and Frechet JMJ (1996) Stimuli-responsive hybrid macromolecules: novel 
amphiphilic star copolymers with dendritic groups at the periphery. / Am Chem Soc 
118:3785-3786. 
40. Liu M, Kono K and Frechet JM (2000) Water-soluble dendritic unimolecular micelles: 
Their potential as drug delivery agents. / Control Rel 65:121-131. 
41. Kojima C, Kono K, Maruyama K and Takagishi T (2000) Synthesis of polyamidoamine 
dendrimers having poly(ethylene glycol) grafts and their ability to encapsulate anticancer 
drugs. Bioconjug Chem 11:910-917. 
42. Wang F, Bronich TK, Kabanov AV, Rauh RD and Roovers J (2005) Synthesis and evaluation 
of a star amphiphilic block copolymer from poly(epsilon-caprolactone) and 
poly(ethylene glycol) as a potential drug delivery carrier. Bioconjug Chem 16:397-405. 
43. Heise A, Hedrick JL, Frank CW and Miller RD (1999) Starlike block copolymers with 
amphiphilic arms as models for unimolecular micelles. / Am Chem Soc 121:8647-8648. 
44. Liu H, Jiang A, Guo J and Uhrich KE (1999) Unimolecular micelles: synthesis and 
characterization of amphiphilic polymer systems. / Polym Sci Part A: Polym Chem 37: 
703-711. 
45. Antoun S, Gohy JF and Jerome R (2001) Micellization of quaternized poly(2- 
(dimethylamino)ethyl methacrylate)-block-poly(methyl methacrylate) copolymers in 
water. Polym 42:3641-3648. 
46. Jones M-C, Ranger M and Leroux J-C (2003) pH-sensitive unimolecular polymeric 
micelles: Synthesis a novel drug carrier. Bioconjug Chem 14:774-781. 
47. Morton M, Helminiak TE, Gadkary SD and Bueche F (1962) Preparation and properties 
of monodisperse branched polystyrene. / Polym Sci 57:471^82. 
48. Gauthier M, Li J and Dockendorff J (2003) Arborescent polystyrene-graft-poly(2- 
vinylpyridine) copolymers as unimolecular micelles. Synthesis from acetylated substrates. 
Macromolecules 36:2642-2648. 
49. Iijima M, Nagasaki Y, Okada T, Kato M and Kataoka K (1999) Core-polymerized 
reactive micelles from heterotelechelic amphiphilic block copolymers. Macromolecules 
32:1140-1146. 
50. Kim J-H, Emoto K, Iijima M, Nagasaki Y, Aoyagi T, Okano T, Sakurai Y and Kataoka K 
(1999) Core-stabilized polymeric micelle as potential drug carrier: increased solubilization 
of taxol. Polym Adv Technol 10:647-654. 
51. Guo A, Liu G and Tao J (1996) Star polymers and nanospheres from cross-linkable 
diblock copolymers. Macromolecules 29:2487-2493. 
52. Won Y-Y, Davis HT and Bates FS (1999) Giant wormlike rubber micelles. Science 
283:960-963. 
53. Rapoport N (1999) Stabilization and activation of Pluronic micelles for tumor-targeted 
drug delivery. Coll SurfB: Biointerf16:93-111. 
54. Thurmond KB, II, Huang H, Clark CG, Jr., Kowalewski T and Wooley KL (1999) Shell 
crosslinked polymer micelles: Stabilized assemblies with great versatility and potential. 
Coll SurfB: Biointerf16:45-54. 
84 Batrakova et al. 
55. Zhang Q, Remsen EE and Wooley KL (2000) Shell cross-linked nanoparticles containing 
hydrolytically degradable, crystalline core domains. / Am Chetn Soc 122: 
3642-3651. 
56. Buetuen V, Lowe AB, Billingham NC and Armes SP (1999) Synthesis of zwitterionic 
shell cross-linked micelles. / Am Chem Soc 121:4288-4289. 
57. Bronich TK and Kabanov AV (2004) Novel block ionomer micelles with cross-linked 
ionic cores. Polym Prepr 45:384-385. 
58. Bader H, Ringsdorf H and Schmidt B (1984) Water-soluble polymers in medicine. Angew 
Makromol Chem 123/124:457-485. 
59. Duncan R (2003) The dawning era of polymer therapeutics. Nat Rev Drug Discov 
2:347-360. 
60. Bulmus V, Woodward M, Lin L, Murthy N, Stayton P and Hoffman A (2003) A 
new pH-responsive and glutathione-reactive, endosomal membrane-disruptive polymeric 
carrier for intracellular delivery of biomolecular drugs. / Control Rel 93: 
105-120. 
61. Veronese FM and Morpurgo M (1999) Bioconjugation in pharmaceutical chemistry. 
Farmaco 54:497-516. 
62. D'Souza AJ and Topp EM (2004) Release from polymeric prodrugs: Linkages and their 
degradation. / Pharm Sci 93:1962-1979. 
63. Bae Y, Fukushima S, Harada A and Kataoka K (2003) Design of environment-sensitive 
supramolecular assemblies for intracellular drug delivery: Polymeric micelles that 
are responsive to intracellular pH change. Angewandte Chemie, Int Edn 42:4640^643, 
S4640/1-S4640/11. 
64. Bae Y, Nishiyama N, Fukushima S, Koyama H, Yasuhiro M and Kataoka K (2005) Preparation 
and biological characterization of polymeric micelle drug carriers with intracellular 
pH-triggered drug release property: Tumor permeability, controlled subcellular 
drug distribution, and enhanced in vivo antitumor efficacy. Bioconjug Chem 16:122-130. 
65. Yokoyama M (1992) Block copolymers as drug carriers. Crit Rev Ther Drug Can Syst 
9:213-248. 
66. Yokoyama M, Sugiyama T, Okano T, Sakurai Y, Naito M and Kataoka K (1993) Analysis 
of micelle formation of an adriamycin-conjugated polyethylene glycol-poly(aspartic 
acid) block copolymer by gel permeation chromatography. Pharm Res 10:895-899. 
67. Yokoyama M, Fukushima S, Uehara R, Okamoto K, Kataoka K, Sakurai Y and Okano T 
(1998) Characterization of physical entrapment and chemical conjugation of adriamycin 
in polymeric micelles and their design for in vivo delivery to a solid tumor. 
/ Control Rel 50:79-92. 
68. Kabanov AV, Chekhonin VP, Alakhov VY, Batrakova EV, Lebedev AS, Melik-Nubarov 
NS, Arzhakov SA, Levashov AV, Morozov GV, Severin ES and Kabanov VA (1989) 
The neuroleptic activity of haloperidol increases after its solubilization in surfactant 
micelles. Micelles as microcontainers for drug targeting. FEBS Lett 258:343-345. 
69. La SB, Okano T and Kataoka K (1996) Preparation and characterization of the micelleforming 
polymeric drug indomethacin-incorporated poly(ethylene oxide)-poly(betabenzyl 
L-aspartate) block copolymer micelles. / Pharm Sci 85:85-90. 
Polymer Micelles as Drug Carriers 85 
70. Yokoyama M, Satoh A, Sakurai Y, Okano T, Matsumura Y, Kakizoe T and Kataoka K 
(1998) Incorporation of water-insoluble anticancer drug into polymeric micelles and 
control of their particle size. / Control Rel 55:219-229. 
71. Inoue T, Chen G, Nakamae K and Hoffman AS (1998) An AB block copolymer of 
oligo(methyl methacrylate) and poly(acrylic acid) for micellar delivery of hydrophobic 
drugs. / Control Rel 51:221-229. 
72. Allen C, Han J, Yu Y, Maysinger D and Eisenberg A (2000) Polycaprolactone-bpoly(
ethylene oxide) copolymer micelles as a delivery vehicle for dihydrotestosterone. 
/ Control Rel 63:275-286. 
73. Kabanov AV, Batrakova EV, Melik-Nubarov NS, Fedoseev NA, Dorodnich TY, Alakhov 
VY, Chekhonin VP, Nazarova IR and Kabanov VA (1992) A new class of drug carriers: 
Micelles of poly(oxyethylene)-poly(oxypropylene) block copolymers as microcontainers 
for drug targeting from blood in brain. / Control Rel 22:141-157. 
74. Burt HM, Zhang X, Toleikis P, Embree L and Hunter WL (1999) Development of copolymers 
of poly(DL-lactide) and methoxypolyethylene glycol as micellar carriers of paclitaxel. 
Coll Surf, B: Biointerf 16:161-171. 
75. Lavasanifar A, Samuel J and Kwon GS (2001) Micelles self-assembled from 
poly(ethylene oxide)-block-poly(N-hexyl stearate L-aspartamide) by a solvent evaporation 
method: Effect on the solubilization and haemolytic activity of amphotericin B. 
/ Control Rel 77:155-160. 
76. Lavasanifar A, Samuel J, Sattari S and Kwon GS (2002) Block copolymer micelles for 
the encapsulation and delivery of amphotericin B. Pharm Res 19:418-^22. 
77. Hurter PN, Scheutjens JMHM and Hatton TA (1993) Molecular modeling of micelle 
formation and solubilization in block copolymer micelles. 1. A self-consistent meanfield 
lattice theory. Macromolecules 26:5592-5601. 
78. Nagarajan R and Ganesh K (1996) Comparison of solubilization of hydrocarbons in 
(PEO-PPO) diblock versus (PEO-PPO-PEO) triblock copolymer micelles. / Coll Interf 
Sci 184:489-499. 
79. Xing L and Mattice WL (1997) Strong solubilization of small molecules by triblockcopolymer 
micelles in selective solvents. Macromolecules 30:1711-1717. 
80. Gadelle F, Koros WJ and Schechter RS (1995) Solubilization of aromatic solutes in block 
copolymers. Macromolecules 28:4883-4892. 
81. Nagarajan R, Barry M and Ruckenstein E (1986) Unusual selectivity in solubilization 
by block copolymer micelles. Langmuir 2:210-215. 
82. Liu J, Xiao Y and Allen C (2004) Polymer-drug compatibility: A guide to the development 
of delivery systems for the anticancer agent, ellipticine. / Pharm Sci 93:132-143. 
83. Nagarajan R and Ganesh K (1989) Block copolymer self-assembly in selective solvents: 
Theory of solubilization in spherical micelles. Macromolecules 22:4312-4325. 
84. Hurter PN and Hatton TA (1992) Solubilization of polycyclic aromatic hydrocarbons 
by poly(ethylene oxide-propylene oxide) block copolymer micelles: Effects of polymer 
structure. Langmuir 8:1291-1299. 
85. Kozlov MY, Melik-Nubarov NS, Batrakova EV and Kabanov AV (2000) Relationship 
between pluronic block copolymer structure, critical micellization concentration 
86 Batrakova etal. 
and partitioning coefficients of low molecular mass solutes. Macromolecules 33: 
3305-3313. 
86. Hurter PN, Scheutjens JMHM and Hatton TA (1993) Molecular modeling of micelle formation 
and solubilization in block copolymer micelles. 2. Lattice theory for monomers 
with internal degrees of freedom. Macromolecules 26:5030-5040. 
87. Teng Y, Morrison ME, Munk P, Webber SE and Prochazka K (1998) Release kinetics 
studies of aromatic molecules into water from block polymer micelles. Macromolecules 
31:3578-3587. 
88. Choucair A and Eisenberg A (2003) Interfacial solubilization of model amphiphilic 
molecules in block copolymer micelles. J Am Chem Soc 125:11993-12000. 
89. Kim SY, Shin IG, Lee YM, Cho CS and Sung YK (1998) Methoxy polyethylene 
glycol) and epsilon-caprolactone amphiphilic block copolymeric micelle containing 
indomethacin. II. Micelle formation and drug release behaviours. / Control Rel 
51:13-22. 
90. Kabanov AV, Vinogradov SV, Suzdaltseva YG and Alakhov VY (1995) Water-soluble 
block polycations as carriers for oligonucleotide delivery. Bioconjug Chem 6:639-643. 
91. Harada A and Kataoka K (1995) Formation of polyion complex micelles in an aqueous 
milieu from a pair of oppositely-charged block copolymers with poly(ethylene glycol) 
segments. Macromolecules 28:5294-5299. 
92. Kabanov VA and Kabanov AV (1998) Interpolyelectrolyte and block ionomer complexes 
for gene delivery: Physico-chemical aspects. Adv Drug Del Rev 30:49-60. 
93. Solomatin SV, Bronich TK, Kabanov VA, Eisenberg A and Kabanov AV (2001) Block 
ionomer complexes: Novel environmentally responsive materials. Polym Prepr 42: 
107-108. 
94. Solomatin SV, Bronich TK, Eisenberg A, Kabanov VA and Kabanov AV (2003) Environmentally 
responsive nanoparticles from block ionomer complexes: Effects of pH and 
ionic strength. Langmuir 19:8069-8076. 
95. Zhang G-D, Harada A, Nishiyama N, Jiang D-L, Koyama H, Aida T and Kataoka K 
(2003) Polyion complex micelles entrapping cationic dendrimer porphyrin: Effective 
photosensitizer for photodynamic therapy of cancer. / Control Rel 93:141-150. 
96. Itaka K, Kanayama N, Nishiyama N, Jang W-D, Yamasaki Y, Nakamura K, 
Kawaguchi H and Kataoka K (2004) Supramolecular Nanocarrier of siRNA from PEGBased 
Block Catiomer Carrying Diamine Side Chain with Distinctive pKa Directed To 
Enhance Intracellular Gene Silencing. / Am Chem Soc 126:13612-13613. 
97. Kabanov AV and Bronich TK (2002) Structure, dispersion stability and dynamics 
of DNA and polycation complexes, in Kim SW and Mahato R (eds.) Pharmaceutical 
Perspectives of Nucleic Acid-Based Therapeutics. Taylor & Francis; London, New York, 
pp. 164-189. 
98. Yokoyama M, Okano T, Sakurai Y, Suwa S and Kataoka K (1996) Introduction of cisplatin 
into polymeric micelle. / Control Rel 39:351-356. 
99. Nishiyama N, Yokoyama M, Aoyagi T, Okano T, Sakurai Y and Kataoka K (1999) 
Preparation and characterization of self-assembled polymer-metal complex micelle 
Polymer Micelles as Drug Carriers 87 
from cis-dichlorodiammineplatinum(II) and poly(ethylene glycol)-poly(alpha.,.beta.- 
aspartic acid) block copolymer in an aqueous medium. Langmuir 15:377-383. 
100. Holleb AI, Fink DJ and Murphy GP (eds.) (1991) American Cancer Society Textbook of 
Clinical Oncology. American Cancer Society: Atlanta, GA. 
101. Sherman SE and Lippard SJ (1987) Structural aspects of platinum anticancer drug interactions 
with DNA. Chem Rev 87:1153-1181. 
102. Nishiyama N and Kataoka K (2001) Preparation and characterization of size-controlled 
polymeric micelle containing cis-dichlorodiammineplatinum(II) in the core. / Control 
Rel 74:83-94. 
103. Nishiyama N, Okazaki S, Cabral H, Miyamoto M, Kato Y, Sugiyama Y, Nishio K, 
Matsumura Y and Kataoka K (2003) Novel cisplatin-incorporated polymeric micelles 
can eradicate solid tumors in mice. Cancer Res 63:8977-8983. 
104. Jang W-D, Nishiyama N, Zhang G-D, Harada A, Jiang D-L, Kawauchi S, Morimoto Y, 
Kikuchi M, Koyama H, Aida T and Kataoka K (2005) Supramolecular nanocarrier of 
anionic dendrimer porphyrins with cationic block copolymers modified with polyethylene 
glycol to enhance intracellular photodynamic efficacy. Angewandte Chemie, Int Edn 
44:419^23. 
105. Stapert HR, Nishiyama N, Jiang D-L, Aida T and Kataoka K (2000) Polyion complex 
micelles encapsulating light-harvesting ionic dendrimer zinc porphyrins. Langmuir 
16:8182-8188. 
106. Harada A and Kataoka K (1998) Novel polyion complex micelles entrapping enzyme 
molecules in the core: Preparation of narrowly-distributed micelles from lysozyme 
and poly(ethylene glycol)-poly(aspartic acid) block copolymer in aqueous medium. 
Macromolecules 31:288-294. 
107. Kawamura A, Yoshioka Y, Harada A and Kono K (2005) Acceleration of enzymatic reaction 
of trypsin through the formation of water-soluble complexes with poly(ethylene 
glycol)-block-poly(a,b-aspartic acid). Biomacromolecules 6:627-631. 
108. Harada A and Kataoka K (2003) Switching by pulse electric field of the elevated enzymatic 
reaction in the core of polyion complex micelles.} Am Chem Soc 125:15306-15307. 
109. Bronich TK, Kabanov AV, Kabanov VA, Yu K and Eisenberg A (1997) Soluble complexes 
from poly(ethylene oxide)-block-polymethacrylate anions and N-alkylpyridinium 
cations. Macromolecules 30:3519-3525. 
110. Bronich TK, Cherry T, Vinogradov SV, Eisenberg A, Kabanov VA and Kabanov AV 
(1998) Self-assmbly in mixtures of poly(ethylene oxide)-graft-poly(ethyleneimine) and 
alkyl sulfates. Langmuir 14:6101-6106. 
111. Kabanov AV, Bronich TK, Kabanov VA, Yu K and Eisenberg A (1998) Spontaneous 
formation of vesicles from complexes of block ionomers and surfactants. / Am Chem 
Soc 120:9941-9942. 
112. Bronich TK, Popov AM, Eisenberg A, Kabanov VA and Kabanov AV (2000) Effects of 
block length and structure of surfactant on self-assembly and solution behavior of block 
ionomer complexes. Langmuir 16:481^89. 
113. Bronich TK, Ouyang M, Kabanov VA, Eisenberg A, Szoka FC, Jr. and Kabanov AV 
(2002) Synthesis of vesicles on polymer template. / Am Chem Soc 124:11872-11873. 
88 Batrakova et al. 
114. Bronich TK, Nehls A, Eisenberg A, Kabanov VA and Kabanov AV (1999) Novel drug 
delivery systems based on the complexes of block ionomers and surfactants of opposite 
charge. Col Surf B 16:243-251. 
115. Solomatin SV, Bronich TK, Eisenberg A, Kabanov VA and Kabanov AV (2004) Colloidal 
stability of aqueous dispersions of block ionomer complexes: Effects of temperature 
and salt. Langmuir 20: 2066-2068. 
116. Pinzani V, Bressolle F, Haug IJ, Galtier M, Blayac JP and Balmes P (1994) Cisplatininduced 
renal toxicity and toxicity-modulating strategies: A review. Cancer Chemother 
Pharmacol 35:1-9. 
117. Weinstein DM, Mihm MJ and Bauer JA (2000) Cardiac peroxynitrite formation and left 
ventricular dysfunction following doxorubicin treatment in mice. / Pharmacol Exp Ther 
294:396-401. 
118. Kwon GS and Kataoka K (1995) Block copolymer micelles as long-circulating drug 
vehicles. Adv Drug Del Rev 16:295-309. 
119. Batrakova EV, Li S, Li Y, Alakhov VY, Elmquist WF and Kabanov AV (2004) Distribution 
kinetics of a micelle-forming block copolymer Pluronic P85. / Control Rel 100:389-397. 
120. Fraser JR, Laurent TC, Pertoft H and Baxter E (1981) Plasma clearance, tissue distribution 
and metabolism of hyaluronic acid injected intravenously in the rabbit. Biochem J 
200:415^124. 
121. Kissel M, Peschke P, Subr V, Ulbrich K, Schuhmacher J, Debus J and Friedrich E 
(2001) Synthetic macromolecular drug carriers: Biodistribution of poly[(N-2- 
hydroxypropyl)methacrylamide] copolymers and their accumulation in solid rat 
tumors. PDAJPharm Sci Technol 55:191-201. 
122. Matsumura Y and Maeda H (1986) A new concept for macromolecular therapeutics in 
cancer chemotherapy: Mechanism of tumoritropic accumulation of proteins and the 
antitumor agent smancs. Cancer Res 46:6387-6392. 
123. Maeda H (2001) The enhanced permeability and retention (EPR) effect in tumor vasculature: 
The key role of tumor-selective macromolecular drug targeting. Adv Enzyme 
Regul 41:189-207. 
124. Giblin MF, Veerendra B and Smith CJ (2005) Radiometallation of receptor-specific peptides 
for diagnosis and treatment of human cancer. In Vivo 19:9-29. 
125. Rini BI (2005) VEGF-targeted therapy in metastatic renal cell carcinoma. Oncologist 
10:191-197. 
126. Lin MZ, Teitell MA and Schiller GJ (2005) The evolution of antibodies into versatile 
tumor-targeting agents. Clin Cancer Res 11:129-138. 
127. Levi FA, Hrushesky WJ, Halberg F, Langevin TR, Haus E and Kennedy BJ (1982) Lethal 
nephrotoxicity and hematologic toxicity of cis-diamminedichloroplatinum ameliorated 
by optimal circadian timing and hydration. Eur } Cancer Clin Oncol 18:471-477. 
128. Greish K, Sawa T, Fang J, Akaike T and Maeda H (2004) SMA-doxorubicin, a new 
polymeric micellar drug for effective targeting to solid tumours. / Control Rel 97: 
219-230. 
129. Lee ES, Na K and Bae YH (2005) Doxorubicin loaded pH-sensitive polymeric micelles 
for reversal of resistant MCF-7 tumor. / Control Rel 103:405-418. 
Polymer Micelles as Drug Carriers 89 
130. Alakhov V, Klinski E, Li S, Pietrzynski G, Venne A, Batrakova E, Bronitch T and Kabanov 
AV (1999) Block copolymer-based formulation of doxorubicin. From cell screen to clinical 
trials. Coll SurfB: Biointerf'16:113-134. 
131. Kim SC, Kim DW, Shim YH, Bang JS, Oh HS, Wan Kim S and Seo MH (2001) In vivo 
evaluation of polymeric micellar paclitaxel formulation: Toxicity and efficacy. / Control 
Rel 72:191-202. 
132. Torchilin VP, Lukyanov AN, Gao Z and Papahadjopoulos-Sternberg B (2003) Immunomicelles: 
Targeted pharmaceutical carriers for poorly soluble drugs. Proc Natl Acad 
Sci USA 100:6039-6044. 
133. Taillefer J, Brasseur N, van Lier JE, Lenaerts V, Le Garrec D and Leroux JC (2001) In vitro 
and in vivo evaluation of pH-responsive polymeric micelles in a photodynamic cancer 
therapy model. / Pharm Pharmacol 53:155-166. 
134. Le Garrec D, Taillefer J, Van Lier JE, Lenaerts V and Leroux JC (2002) Optimizing pHresponsive 
polymeric micelles for drug delivery in a cancer photodynamic therapy 
model. / Drug Targ 10:429-437. 
135. Yokoyama M, Okano T, Sakurai Y, Fukushima S, Okamoto K and Kataoka K (1999) 
Selective delivery of adriamycin to a solid tumor using a polymeric micelle carrier 
system. / Drug Targ 7:171-186. 
136. Greish K, Nagamitsu A, Fang J and Maeda H (2005) Copoly(styrene-maleic acid)- 
pirarubicin micelles: High tumor-targeting efficiency with little toxicity. Bioconjug Chem 
16:230-236. 
137. Nasongkla N, Shuai X, Ai H, Weinberg BD, Pink J, Boothman DA and Gao J (2004) 
cRGD-functionalized polymer micelles for targeted doxorubicin delivery. Angew Chem 
Int Ed Engl 43:6323-6327. 
138. Paulos CM, Turk MJ, Breur GJ and Low PS (2004) Folate receptor-mediated targeting 
of therapeutic and imaging agents to activated macrophages in rheumatoid arthritis. 
Adv Drug Del Rev 56:1205-1217. 
139. Yoo HS and Park TG (2004) Folate receptor targeted biodegradable polymeric doxorubicin 
micelles. / Control Rel 96:273-283. 
140. Allen TM (2002) Ligand-targeted therapeutics in anticancer therapy. Nat Rev Cancer 
2:750-763. 
141. Gao Z, Lukyanov AN, Chakilam AR and Torchilin VP (2003) PEG-PE/ phosphatidylcholine 
mixed immunomicelles specifically deliver encapsulated taxol to tumor cells 
of different origin and promote their efficient killing. / Drug Targ 11:87-92. 
142. Tannock IF and Rotin D (1989) Acid pH in tumors and its potential for therapeutic 
exploitation. Cancer Res 49:4373^1384. 
143. Kataoka K, Matsumoto T, Yokoyama M, Okano T, Sakurai Y, Fukushima S, Okamoto K 
and Kwon GS (2000) Doxorubicin-loaded poly(ethylene glycol)-poly(beta-benzyl-Laspartate) 
copolymer micelles: Their pharmaceutical characteristics and biological significance. 
/ Control Rel 64:143-153. 
144. Yoo HS, Lee EA and Park TG (2002) Doxorubicin-conjugated biodegradable polymeric 
micelles having acid-cleavable linkages. / Control Rel 82:17-27. 
90 Batrakova et al. 
145. Hruby M, Konak C and Ulbrich K (2005) Polymeric micellar pH-sensitive drug delivery 
system for doxorubicin. / Control Rel 103:137-148. 
146. Kabanov AV, Bronich TK, Kabanov VA, Yu K and Eisenberg A (1996) Soluble stoichiometric 
complexes from poly(N-ethyl-4-vinylpyridinium) cations and poly(ethylene 
oxide)-Wocfc-polymethacrylate anions. Macromolecules 29:6797-6802. 
147. Lee ES, Shin HJ, Na K and Bae YH (2003) Poly(L-histidine)-PEG block copolymer 
micelles and pH-induced destabilization. / Control Rel 90:363-374. 
148. Alakhov VY, Moskaleva EY, Batrakova EV and Kabanov AV (1996) Hypersensitization 
of multidrug resistant human ovarian carcinoma cells by pluronic P85 block copolymer. 
Bioconjug Chem 7:209-216. 
149. Venne A, Li S, Mandeville R, Kabanov Aand Alakhov V (1996) Hypersensitizing effect of 
pluronic L61 on cytotoxic activity, transport, and subcellular distribution of doxorubicin 
in multiple drug- resistant cells. Cancer Res 56:3626-3629. 
150. Batrakova EV, Dorodnych TY, Klinskii EY, Kliushnenkova EN, Shemchukova OB, 
Goncharova ON, Arjakov SA, Alakhov VY and Kabanov AV (1996) Anthracycline 
antibiotics non-covalently incorporated into the block copolymer micelles: In vivo evaluation 
of anti cancer activity. Br } Cancer 74:1545-1552. 
151. Kabanov AV, Batrakova EV and Alakhov VY (2002) Pluronic block copolymers for 
overcoming drug resistance in cancer. Adv Drug Deliv Rev 54:759-779. 
152. Batrakova EV, Lee S, Li S, Venne A, Alakhov V and Kabanov A (1999) Fundamental 
relationships between the composition of pluronic block copolymers and their hypersensitization 
effect in MDR cancer cells. Pharm Res 16:1373-1379. 
153. Miller DW, Batrakova EV and Kabanov AV (1999) Inhibition of multidrug resistanceassociated 
protein (MRP) functional activity with pluronic block copolymers. Pharm 
Res 16:396-401. 
154. Batrakova EV, Li S, Alakhov VY, Elmquist WF, Miller DW and Kabanov AV (2003) Sensitization 
of cells overexpressing multidrug-resistant proteins by pluronic P85. Pharm 
Res 20:1581-1590. 
155. Minko T, Batrakova E, Li S, Li Y, Pakunlu R, Alakhov V and Kabanov A (2005) Pluronic 
block copolymers alter apoptotic signal transduction of doxorubicin in drug-resistant 
cancer cells. / Control Rel. 
156. Batrakova EV, Han HY, Miller DW and Kabanov AV (1998) Effects of pluronic P85 
unimers and micelles on drug permeability in polarized BBMEC and Caco-2 cells. 
Pharm Res 15:1525-1532. 
157. Evers R, Kool M, Smith AJ, van Deemter L, de Haas M and Borst P (2000) Inhibitory 
effect of the reversal agents V-104, GF120918 and Pluronic L61 on MDR1 Pgp-, MRP1- 
and MRP2-mediated transport. Br J Cancer 83:366-374. 
158. Batrakova EV, Miller DW, Li S, Alakhov VY, Kabanov AV and Elmquist WF (2001) 
Pluronic P85 enhances the delivery of digoxin to the brain: In vitro and in vivo studies. 
/ Pharmacol Exp Ther 296:551-557. 
159. Miller DW, Batrakova EV, Waltner TO, Alakhov V and Kabanov AV (1997) Interactions 
of pluronic block copolymers with brain microvessel endothelial cells: Evidence of two 
potential pathways for drug absorption. Bioconjug Chem 8:649-657. 
Polymer Micelles as Drug Carriers 91 
160. Batrakova EV, Li S, Elmquist WF, Miller DW, Alakhov VY and Kabanov AV (2001) 
Mechanism of sensitization of MDR cancer cells by Pluronic block copolymers: Selective 
energy depletion. Br J Cancer 85:1987-1997. 
161. Batrakova EV, Li S, Vinogradov SV, Alakhov VY, Miller DW and Kabanov AV (2001) 
Mechanism of pluronic effect on P-glycoprotein efflux system in blood-brain barrier: 
Contributions of energy depletion and membrane fluidization. / Pharmacol Exp Ther 
299:483-493. 
162. Batrakova EV, Li S, Alakhov VY, Miller DW and Kabanov AV (2003) Optimal structure 
requirements for Pluronic block copolymers in modifying P-glycoprotein drug efflux 
transporter activity in bovine brain microvessel endothelial cells. / Pharmacol Exp Ther 
304:845-854. 
163. Rapoport N, Marin AP and Timoshin AA (2000) Effect of a polymeric surfactant on 
electron transport in HL-60 cells. Arch Biochem Biophys 384:100-108. 
164. Kabanov AV, Batrakova EV and Alakhov VY (2003) An essential relationship between 
ATP depletion and chemosensitizing activity of Pluronic block copolymers. / Control 
Rel 91:75-83. 
165. Batrakova EV, Li S, Miller DW and Kabanov AV (1999) Pluronic P85 increases permeability 
of a broad spectrum of drugs in polarized BBMEC and Caco-2 cell monolayers. 
Pharm Res 16:1366-1372. 
166. Batrakova EV, Zhang Y, Li Y, Li S, Vinogradov SV, Persidsky Y, Alakhov V, Miller DW 
and Kabanov AV (2004) Effects of Pluronic P85 on GLUT1 and MCT1 transporters in 
the blood brain barrier. Pharm Res in press. 
167. Kabanov AV, Batrakova EV and Miller DW (2003) Pluronic((R)) block copolymers as 
modulators of drug efflux transporter activity in the blood-brain barrier. Adv Drug Del 
Rev 55:151-164. 
168. Kabanov AV and Batrakova EV (2004) New technologies for drug delivery across the 
blood brain barrier. Curr Pharm Des 10:1355-1363. 
169. Kwon GS (2003) Polymeric micelles for delivery of poorly water-soluble compounds. 
Crit Rev Ther Drug Carr Syst 20:357-403. 
170. Adams ML and Kwon GS (2003) Relative aggregation state and hemolytic activity 
of amphotericin B encapsulated by poly(ethylene oxide)-block-poly(N-hexyl- 
L-aspartamide)-acyl conjugate micelles: Effects of acyl chain length. / Control Rel 
87:23-32. 
171. Adams ML, Andes DR and Kwon GS (2003) Amphotericin B encapsulated in micelles 
based on poly(ethylene oxide)-block-poly(L-amino acid) derivatives exerts reduced 
in vitro hemolysis but maintains potent in vivo antifungal activity. Biomacromolecules 
4:750-757. 
172. Croy SR and Kwon GS (2004) The effects of Pluronic block copolymers on the aggregation 
state of nystatin. / Control Rel 95:161-171. 
173. Jagannath C, Sepulveda E, Actor JK, Luxem F, Emanuele MR and Hunter RL 
(2000) Effect of poloxamer CRL-1072 on drug uptake and nitric-oxide-mediated 
killing of Mycobacterium avium by macrophages. Immunopharmacology 48: 
185-197. 
92 Batrakova et al. 
174. Jagannath C, Emanuele MR and Hunter RL (2000) Activity of poloxamer CRL- 
1072 against drug-sensitive and resistant strains of Mycobacterium tuberculosis in 
macrophages and in mice. Int J Antimicrob Agents 15:55-63. 
175. Jagannath C, Emanuele MR and Hunter RL (1999) Activities of poloxamer CRL-1072 
against Mycobacterium avium in macrophage culture and in mice. Antimicrob Agents 
Chemother 43:2898-2903. 
176. Jagannath C, Wells A, Mshvildadze M, Olsen M, Sepulveda E, Emanuele M, Hunter 
RL, Jr. and Dasgupta A (1999) Significantly improved oral uptake of amikacin in FVB 
mice in the presence of CRL-1605 copolymer. Life Sci 64:1733-1738. 
177. Torchilin VP (2002) PEG-based micelles as carriers of contrast agents for different imaging 
modalities. Adv Drug Del Rev 54:235-252. 
178. Trubetskoy VS, Frank-Kamenetsky MD, Whiteman KR, Wolf GL and Torchilin VP (1996) 
Stable polymeric micelles: Lymphangiographic contrast media for gamma scintigraphy 
and magnetic resonance imaging. Acad Radiol 3:232-238. 
179. Trubetskoy VS, Gazelle GS, Wolf GL and Torchilin VP (1997) Block-copolymer of 
polyethylene glycol and polylysine as a carrier of organic iodine: Design of longcirculating 
particulate contrast medium for X-ray computed tomography. / Drug Targ 
4:381-388. 
180. Weissig V, Whiteman KR and Torchilin VP (1998) Accumulation of protein-loaded longcirculating 
micelles and liposomes in subcutaneous Lewis lung carcinoma in mice. 
Pharm Res 15:1552-1556. 
181. Katayose S and Kataoka K (1997) Water-soluble polyion complex associates of DNA 
and poly(ethylene glycol)-poly(L-lysine) block copolymer. Bioconj Chem 8:702-707. 
182. Wolfert MA, Schacht EH, Toncheva V, Ulbrich K, Nazarova O and Seymour LW (1996) 
Characterization of vectors for gene therapy formed by self-assembly of DNA with 
synthetic block co-polymers. Hum Gene Ther 7:2123-2133. 
183. Vinogradov SV, Bronich TK and Kabanov AV (1998) Self-assembly of polyaminepoly(
ethylene glycol) copolymers with phosphorothioate oligonucleotides. Bioconjug 
Chem 9:805-812. 
184. Itaka K, Harada A, Nakamura K, Kawaguchi H and Kataoka K (2002) Evaluation by 
fluorescence resonance energy transfer of the stability of nonviral gene delivery vectors 
under physiological conditions. Biomacromolecules 3:841-845. 
185. Roy S, Zhang K, Roth T, Vinogradov S, Kao RS and Kabanov A (1999) Reduction of 
fibronectin expression by intravitreal administration of antisense oligonucleotides. Nat 
Biotechnol 17:476-479. 
186. Ogris M, Steinlein P, Kursa M, Mechtler K, Kircheis R and Wagner E (1998) The size of 
DNA/transferrin-PEI complexes is an important factor for gene expression in cultured 
cells. Gene Ther 5:1425-1433. 
187. Oupicky D, Ogris M, Howard KA, Dash PR, Ulbrich K and Seymour LW (2002) Importance 
of lateral and steric stabilization of polyelectrolyte gene delivery vectors for 
extended systemic circulation. Mol Ther 5:463-472. 
Polymer Micelles as Drug Carriers 93 
188. Harada-Shiba M, Yamauchi K, Harada A, Takamisawa I, Shimokado K and Kataoka K 
(2002) Polyion complex micelles as vectors in gene therapy-pharmacokinetics and 
in vivo gene transfer. Gene Ther 9:407-414. 
189. Choi YH, Liu F> ParkJS and Kim SW (1998) Lactose-poly(ethylene glycol)-grafted poly- 
L-lysine as hepatoma cell- tapgeted gene carrier. Bioconjug Chem 9:708-718. 
190. Vinogradov S, Batrakova E, Li S and Kabanov A (1999) Polyion complex micelles with 
protein-modified corona for receptor-mediated delivery of oligonucleotides into cells. 
Bioconjug Chem 10:851-860. 
191. Ward CM, Pechar M, Oupicky D, Ulbrich K and Seymour LW (2002) Modification of 
pLL/DNA complexes with a multivalent hydrophilic polymer permits folate-mediated 
targeting in vitro and prolonged plasma circulation in vivo. } Gene Med 4:536-547. 
192. Nguyen HK, Lemieux P, Vinogradov SV, Gebhart CL, Guerin N, Paradis G, Bronich 
TK, Alakhov VY and Kabanov AV (2000) Evaluation of polyether-polyethyleneimine 
graft copolymers as gene transfer agents. Gene Ther 7:126-138. 
193. Gebhart CL, Sriadibhatla S, Vinogradov S, Lemieux P, Alakhov V and Kabanov AV 
(2002) Design and formulation of polyplexes based on pluronic-polyethyleneimine 
conjugates for gene transfer. Bioconjug Chem 13:937-944. 
194. Belenkov AI, Alakhov VY, Kabanov AV, Vinogradov SV, Panasci LC, Monia BP and 
Chow TY (2004) Polyethyleneimine grafted with pluronic P85 enhances Ku86 antisense 
delivery and the ionizing radiation treatment efficacy in vivo. Gene Ther 11:1665-1672. 
195. Yokoyama M, Miyauchi M, Yamada N, Okano T, Sakurai Y, Kataoka K and Inoue S (1990) 
Characterization and anticancer activity of the micelle-forming polymeric anticancer 
drug adriamycin-conjugated poly(ethylene glycol)-poly(aspartic acid) block copolymer. 
Cancer Res 50:1693-1700. 
196. Nakanishi T, Fukushima S, Okamoto K, Suzuki M, Matsumura Y, Yokoyama M, 
Okano T, Sakurai Y and Kataoka K (2001) Development of the polymer micelle carrier 
system for doxorubicin. / Control Rel 74:295-302. 
197. Danson S, Ferry D, Alakhov V, Margison J, Kerr D, Jowle D, Brampton M, Halbert G 
and Ranson M (2004) Phase I dose escalation and pharmacokinetic study of pluronic 
polymer-bound doxorubicin (SP1049C) in patients with advanced cancer. Br } Cancer 
90:2085-2091. 
198. Valle JW, Lawrance J, Brewer J, Clayton A, Corrie P, Alakhov V and Ranson M (2004) 
A phase II, window study of SP1049C as first-line therapy in inoperable metastatic 
adenocarcinoma of the oesophagus. 2004 ASCO Annual Meeting Vol. Abstract No: 4195. 
199. Kim TY, Kim DW, Chung JY, Shin SG, Kim SC, Heo DS, Kim NK and Bang YJ 
(2004) Phase I and pharmacokinetic study of Genexol-PM, a cremophor-free, polymeric 
micelle-formulated paclitaxel, in patients with advanced malignancies. Clin Cancer Res 
10:3708-3716. 
This page is intentionally left blank
6 
Vesicles Prepared from Synthetic 
Amphiphiles — Polymeric Vesicles and 
Niosomes 
Ijeoma Florence Uchegbu and Andreas G. Schatzlein 
1. Introduction 
This chapter will examine what is known about vesicles prepared from synthetic 
amphiphiles and will encompass a review of the data published on polymeric vesicles 
and non-ionic surfactant vesicles (niosomes). Schematic representations of the 
molecular arrangements in these systems are as depicted in Fig. 1. Examples of 
drug delivery applications will also be presented. 
Vesicular systems arise when amphiphilic molecules self assemble in aqueous 
media in an effort to reduce the high energy interaction between the hydrophobic 
portion of the amphiphile and the aqueous disperse phase, and maximize the low 
energy interaction between the hydrophilic head group and the disperse phase 
(Fig. 1). These self assemblies reside in the nanometre to micrometre size domain. 
Excellent reviews exist on the self assembly of amphiphiles/16 and hence this topic 
will not be dealt with in great detail here. Vesicles are important pharmaceutical 
systems, especially as liposomes, the result of phospholipid self assembly,19 are 
licensed for the clinical delivery of anti cancer drugs.21 It is thus possible that the 
vesicles described here may be incorporated into licensed medicines at some point 
in future. 
95 
96 Uchegbu & Schatzlein 
t % 
If 111 
II 
mnme&m 
Self assembling 
polymerisable monomers 
Polymerisation! 
(a) Polymerised 
vesicles 
mm 
isfi? M# 
V 
* » » * ^ 
*^*^* 
i«s m 
it 
(b) Self assembling 
amphiphilic polymers 
Q 
U 
80% favors dense 
nanoparticles, while a polydactic acid) fraction of 58-80% favors bilayer vesicle 
assemblies, and a polydactic acid) fraction of less than 50% favors the production 
of micellar self assemblies.31 
The sizes of the vesicle and dense nanoparticle assemblies formed from 
amphiphilic poly(ethylenimines) are also dependent on polymer levels of 
hydrophobic modification (mole % cetylation) and the relationships shown in 
Eqs. (1) and (2) have been developed,18 
dv = 1.95Ct + 139 (1) 
dn = 2.31Ct + 5.6 (2) 
where dv = vesicle z-average mean hydrodynamic diameter, Ct = mole% cetylation 
(number of cetyl groups per 100 monomer units), and dn — nanoparticle 
z-average mean hydrodynamic diameter. 
The molecular weight of the polymer is also an important factor to consider 
when choosing vesicle forming polymers. The importance of this parameter has 
been demonstrated with the poly(L-lysine) vesicle system20 [e.g. Compound 6, 
Fig. 3(a)]. With these amphiphiles a vesicle formation index (F') has been computed: 
F' = - ^ = (3) 
LVDP 
where H = mole% unreacted L-lysine units, L — mole% L-lysine units substituted 
with palmitic acid and DP = the degree of polymerisation of the polymer. An F' 
value in excess of 0.168 is necessary for vesicle formation.20 
Additionally, not only does the molecular weight of the polymer impact on 
vesicle formation, but it is also a direct controller of the vesicle mean size; the 
relationship shown in Eq. (4) has been developed for the palmitoyl glycol chitosan 
system,11 
VMW = 0.782dv + 107 (4) 
Vesicles Prepared from Synthetic Amphiphiles 101 
where MW = polymer molecular weight, and dv = vesicle z-average mean 
hydrodynamic diameter. 
2.3. Block copolymers 
Block copolymer vesicles, termed "polymersomes" are fairly new discoveries, being 
first reported in the 1990s.32 Polymersomes have been prepared from a variety of 
block copolymers, some examples of which are given in Fig. 4. There is a clear 
relationship between the hydrophobic content of polymers and self assembly. Low 
levels of hydrophobicity (less than 50% of the polymer consisting of hydrophobic 
HO, X N^- 
J5"H 
HN' 
9 
Fig. 4. Examples of some vesicle forming block copolymers Compound 7,1 Compound 8,7 
and Compound 9.13 
102 Uchegbu & Schatzlein 
moieties) favors the formation of micelles33 and intermediate levels of hydrophobicity 
(50-80%) favors the formation of bilayer vesicles.31,33'34 For the self assembly 
of block copolymers, it has been established that generally the critical packing 
parameter (CPP): 
CPP = ^ (5) 
al 
should approach unity for vesicular self assemblies to prevail,24 where v = volume 
of the hydrophobic block, 1 — length of the hydrophobic block and a = the area of 
the hydrophilic block. 
Vesicle sizes are varied and range from tens of nanometres35 to tens of 
microns.36 Polymersome membranes are 8-21 nm thick; 2-5 times thicker than 
the 4nm membrane thickness displayed by conventional low molecular weight 
amphiphiles.16'27,31'34,35 The thickness of the membrane is determined by the degree 
of polymerization in the hydrophobic block34 and these extra thick membranes 
confer, on the vesicle, exceptional stability to soluble surfactantS24 and mechanical 
stress.24'27,37'38 With these vesicles, there is an asymmetric distribution of the 
polymers in the inner and outer leaflets of the bilayer and polymers with a large 
hydrophilic chain length are preferentially localized to the exterior leaflet and vice 
versa.39 Preferred residence in the outer leaflet is favored by the more hydrophilic 
polymers, because the greater repulsion between the longer hydrophilic corona 
molecules on the outer leaflet stabilize the vesicle curvature.39 
Vesicle stability is a desirable characteristic for pharmaceutical vesicles and as 
such, a great deal of effort has been expended on producing stable systems. As the 
drive for nanomedicines (medicines incorporating functional nanoparticles) grows, 
stability issues will need to be adequately addressed to ensure the widespread 
adoption of such systems. In actual fact, the early workers in the polymeric vesicle 
field were primarily driven by this need to produce stable drug carriers. Extremely 
stable systems are possible on polymerization of block copolymers subsequent to 
self assembly. Poly(ethylene oxide)-WocA:-poly[3-(trimethoxysilyl)propyl methacrylate] 
copolymer vesicles in water, methanol, triethylamine mixtures produced 
polymerized polymersomes that are stable for up to one year.40 Triethylamine 
hydrolyzes the trimethoxysilyl groups and then catalyzes their polycondensation 
to yield an extremely stable hydrophobic polysilsesquioxane core.40,41 Additionally, 
poly(ethylene oxide)-Wocfc-poly(butadiene) vesicles on cross linking produce 
vesicles which are organic solvent resistant.42 
2.4. Preparing vesicles from self-assembling polymers 
Polymeric vesicles are relatively simple to prepare. The input of energy is achieved 
in the laboratory by probe sonication of the amphiphilic polymer in the disperse 
Vesicles Prepared from Synthetic Amphiphiles 103 
phase.1120 However, clearly the energy required for self assembly is not trivial as 
vesicles are not easily formed by hand shaking, unlike low molecular weight surfactant 
formulations.4 Vesicles once formed are morphologically stable for months11 
and may be loaded with hydrophilic43-45 and hydrophobic [see Fig. 6(b) below] 
solutes, by probe sonicating in the presence of such solutes. Commercially, it is 
envisaged that polymeric vesicles may be fabricated by microfluidization and high 
pressure homogenization techniques. 
2.5. Self assembling polymerizable monomers 
Polymerized vesicles may also be prepared by utilizing self assembling polymerizable 
amphiphiles, followed by the polymerization of the resulting vesicular self 
assembly (Fig. 1). Examples of some polymerizable vesicle forming monomers are 
shown in Fig. 5. This method of producing polymerised vesicles is the oldest form 
of polymeric vesicle technology.12,46 
HO-P-OH 
l O 
HO-P-OH 
O 
13 
Fig. 5. Polymerizable vesicle forming monomers used to make polymerized vesicles 
by Jung and others (Compound 10),5 Cho and others (Compound ll),8 Hub and others 
(Compound 12)12 and Bader and others (Compound 13).15 
104 Uchegbu & Schatzlein 
Polymerized vesicles prepared using polymerized self assembling monomers 
are essentially polymer shells and it is unclear how much of the bilayer assembly 
actually survives the polymerization step. The advantage, however, is that 
they are extremely stable, resisting degradation by detergents47-49 or organic 
solvents.8'48,50,51 They are also less leaky,50 thermostable,52 and because the vesicle 
forming components are kinetically trapped by the polymerization process, they 
have improved colloidal stability.8 A major advantage of these nanosystems is that 
they may be isolated as dry powders which are readily dispersible in water to give 
50-100 nm particles;48 thus potentially allowing the formulation of solid vesicle 
dosage forms. Polymerization involves fairly reactive species and hence vesicles 
are best prepared prior to drug loading, which may be a limitation. 
3. Polymeric Vesicle Drug Delivery Applications 
Polymeric vesicles, which are the focus of this chapter, exist in two main varieties as 
illustrated in Fig. 1. These technologies are suitable candidates for the development 
of robust, controllable and responsive nanomedicine drug carriers. 
3.1. Drug targeting 
Poly(oxyethylene) amphiphiles, when incorporated into liposomal26 and niosomal6 
bilayers, prolong vesicle circulation and facilitate tumor targeting,6'53 due to the 
leaky nature of the poorly developed tumor vascular endothelium.54 Only 10 
mole % poly(ethylene oxide) — lipid amphiphiles may be incorporated into 
liposomes55 or niosomes,56'57 without a loss of vesicle integrity due to the preferred 
tendency of the hydrophilic poly(oxyethylene) amphiphiles to form micelles. Polymersomes 
composed of poly(ethylene oxide)-Wocfc-polybutadiene or poly(ethylene 
oxide)-Wocfc-poly(ethylethylene), in which the entire vesicle surface is covered with 
the poly(ethylene oxide) coat, have been studied as long circulating nanocarriers 
for drug delivery58 The circulation time of poly(ethylene oxide) polymersomes is 
directly dependent on the length of the poly(ethylene oxide) block and polymersome 
half lives of up to 28 hrs have been recorded in rats with a poly(ethylene oxide) 
degree of polymerization of 50.58 This half life compares favorably with a half life 
of 14 hrs recorded for poly(oxyethylene) coated liposomes.59 It is assumed that 
the 100% surface coverage of the polymeric vesicles is responsible for the reduced 
clearance of these polymersomes from the blood.38 The long half life of these polymersomes 
makes them excellent candidates for the development of anti tumor 
medicines. 
Furthermore, drug release may be controlled in the polymersomes by controlling 
the hydrolysis rate of the hydrophobic blocks.31 This has been demonstrated 
Vesicles Prepared from Synthetic Amphiphiles 105 
with poly(L-lactic acid)-fr/ocfc-poly(ethylene glycol) and poly(caprolactone)-Wod> 
poly(ethylene glycol) vesicles.31 Hydrolysis of the hydrophobic block causes the 
polymer to move from a vesicular to a micellar assembly, as the overall level of 
hydrophobic content diminishes, and this in turn leads to drug release.31 Hydrolysis 
rates and implicitly release rates may be controlled by varying the relative level 
of the hydrophobic blocks. 
Carbohydrate polymeric vesicles may also be used as drug targeting agents. 
Vesicles prepared from glycol chitosan vesicles improve the intracellular delivery 
of hydrophilic macromolecules44 and anti cancer drugs,45 the latter is achieved with 
the help of a transferrin ligand attached to the surface of the vesicle. 
3.2. Gene delivery 
Poly(L-lysine) based vesicles, prepared from Compound 6 [Fig. 3(a)] have been used 
for gene delivery,29,60 as these vesicles are less toxic than unmodified poly(L-lysine) 
and produce higher levels of gene transfer (Table l).29 The production of polymeric 
vesicles and the resultant reduction in cytotoxicity enables poly(L-lysine) to be used 
in in vivo gene, as the unmodified polymer is too toxic for in vivo use. When the 
targeting ligand, galactose, was bound to the distal ends of the poly(oxyethylene) 
chains, gene expression was increased in HepG2 cells in vitro.60 However, in vivo 
targeting to the liver hepatocytes was not achieved with these systems.60 
A similar procedure with the poly(ethylenimine) vesicles prepared using 
Compound 5 [Fig. 3(a)] also resulted in a reduction in the cytotoxicity of the polymer 
(Table l),17 although in this case, the poly(ethylenimine) vesicles were not as 
efficient gene transfer agents as the free polymer. 
Table 1 Biological Activity of poly(ethylenimine)17 and poly(L-lysine)29 Vesicles. 
Polymer A431 cells A549 
IC50 Gene Transfer IC50 Gene Transfer 
(AtgmL-1) Relative to Parent (jiigmL-1) Relative to Parent 
Polymer Polymer 
Poly(ethylenimine) 1.9 1 5.2 1 
Polymer 5 (Fig. 6(a)) 16.9 0.2 12.6 0.08 
Polymer 5, cholesterol 15.9 0.2 11 0.08 
vesicles 2:1 (gg_1) 
Poly(L-lysine) 7 1 7 1 
Polymer 6 (Fig. 6(a)) 74 7.8 63 2.3 
106 Uchegbu & Schatzlein 
3.3. Responsive release 
The ultimate goal of all drug delivery efforts is the simple fabrication of responsive 
systems that are capable of delivering precise quantities of their pay load in response 
to physiological or more commonly pathological stimuli. Pre-programmable pills, 
implants and injectables are so far merely the unobtainable ideal, however, polymeric 
systems have been fabricated with responsive capability and it is possible 
that in the future, these may be fine tuned to produce truly intelligent and dynamic 
drug delivery devices or systems. 
The various environmental stimuli that may be used to trigger the release of 
encapsulated drug are outlined below and examples are given of existing developments 
in the area. However, in addition to the areas covered below, it may 
be possible in future for pathology specific molecules to interact with polymeric 
vesicles to trigger release. 
3.3.1. pH 
Diblock polypeptides, in which the hydrophilic block consists of ethylene glycol 
derivatised amino acids (L-lysine), and the hydrophobic block consists of poly 
(L-leucine), form pH responsive vesicles which disaggregate at low pH, providing 
the level of L-leucine and polymer chain length is maintained within defined limits 
of about 12-25 mole% and the polymer has a degree of polymerization of less than 
200.13 These L-lysine based systems may be applied to facilitate endosome specific 
release. 
3.3.2. Enzymatic 
Vesicles which release their contents in the presence of an enzyme may be formed 
by loading polymeric vesicles with an enzyme activated prodrug (Fig. 6). The 
particulate nature of the drug delivery system should allow the drug to accumulate 
in tumors, for example, where it may then be activated by an externally 
applied enzyme in a similar manner to the antibody directed enzyme prodrug 
therapeutic strategy. The antibody directed enzyme prodrug therapeutic strategy 
enables an enzyme to be homed to tumors using antibodies followed by the 
application of an enzyme activated prodrug.61 Alternatively, a membrane bound 
enzyme may be used to control and ultimately prolong the activity of either an 
entrapped hydrophilic drug (entrapped in the vesicle aqueous core) or an entrapped 
hydrophobic drug (entrapped in the vesicle membrane) as illustrated in Fig. 6. It is 
possible that the enzyme may be chosen such that it is activated in the presence of 
pathology specific molecules, thus achieving pathology responsive and localized 
drug activity. 
Vesicles Prepared from Synthetic Amphiphiles 107 
2.5 
2.0 
ii 
CD 
1.5 
1.0' 
<2 0.5- 
0.0- 
I 1 
NM/ 
-O— vesicle bound enzyme + external substrate 
- •— external enzyme + vesicle loaded substrate 
-A— control solution + substrate 
--? ^""V 
• • • • • 
W»>* 
0 20 40 60 80 A 
• Enzy m e Time(min) ^ W 
A Water soluble Substrate 
Fig. 6(a). Enzyme activated polymeric vesicles. Vesicles bearing membrane bound enzyme 
(i) were formed by probe sonicating Compound 2 [Fig. 3(a)], cholesterol, N-biotinylated 
dipalmitoyl phosphatidyl ethanolamine (8: 4: 1 gg"1) in neutral phosphate buffer (2mL), 
isolation of the vesicles by ultracentrifugation (150,000 g), redispersion in a similar volume 
of neutral phosphate buffer and incubation of the vesicles with /S-galactosidase streptavidin 
(3 U). Membrane bound enzyme (0.2 mL) was then incubated with o-nitrophenyl-/J-Dgalactoside 
(2.1 mM, 2 mL) and the absorbance monitored (X = 410 ran). The control solution 
contained similar levels of substrate (o-nitrophenyl-jS-D-galactoside) but no enzyme. Vesicles 
encapsulating O-nitrophenyl-jS-D-galactoside (ii) were prepared by probe sonicating Compound 
2, cholesterol (8: 4gg_1) in the presence of o-nitrophenyl-jS-D-galactoside solution 
(34 mM, 2 mL) and isolation of the vesicles by ultracentrifugation and redispersion in neutral 
phosphate buffer. These latter vesicles (0.4 mL) were then incubated with /J-D-galactosidase 
(2UmL_1, 0.1 mL) and the absorbance once again monitored. 
3.3.3. Magnetic 
Magnetically responsive polymerized liposomes composed of 1,2-di (2,4- 
octadecadienoyl)-sn-glycerol-3-phosphorylcholine, loaded with ferric oxide and 
subsequently polymerized may be localized by an external magnetic field to the 
small intestine, and specifically the Payer's patches.47 These polymerized vesicles 
are stable to the degradative influence of solubilizing surfactants such as triton-X 
100,47 and hence should not suffer excessive bile salt mediated degradation during 
gut transit. These magnetically responsive polymeric vesicles improve the absorption 
of drugs via the oral route.47 
108 Uchegbu & Schatzlein 
0.0 10 20 30 40 50.0 
MIN 
^k Membrane bound enzyme 
^ ^ Hydrophobic substrate 
Fig. 6(b). Enzyme activated polymeric vesicles. Vesicles bearing membrane bound 
enzyme and containing the hydrophobic substrate fluorescein di-/S-D-galactospyranoside 
were formed by probe sonicating Compound 2 [Fig. 3(a)], cholesterol, N-biotinylated 
dipalmitoyl phosphatidyl ethanolamine, fluorescein di-^-D-galactospyranoside (8: 4: 1: 
0.0005 g g_1) in neutral phosphate buffer (2 mL) and incubation of the resulting vesicles with 
b-galactosidase streptavidin (0.3 U). The fluorescence of the enzyme hydrolysed substrate 
was then monitored (Excitation wavelength = 490 nm, Emission wavelength = 514 nm). 
3.3.4. Oxygen 
Block copolymer vesicles which are destabilized by oxidative mechanisms have 
been constructed from poly(oxyethylene)-Wocfc-poly(propylene sulphide)-fr/ocfcpoly(
oxyethylene) ABA block copolymers.62 These polymeric vesicles are destabilized 
on the oxidation of the central sulphide block to give sulphoxides and 
ultimately sulphones.62 On oxidation, vesicles are transformed to worm-like 
micelles and finally to spherical micelles, eventually releasing their contents. 
4. Non-ionic Surfactant Vesicles (Niosomes) 
4.1. Self assembly 
The self assembly of non-ionic surfactants into niosomes is dependent on the 
hydrophilic — hydrophobic balance of the surfactant and a CPP (Eq. 1) of between 
0.5-1016 enables niosomal self assembly. Some examples of niosome forming 
molecules are given in Fig. 7. Further molecular specifics that govern niosome 
Vesicles Prepared from Synthetic Amphiphiles 109 
15 
16 
17 
OH HO 
0 /k0J 
- O — v ^ ^ O - 
OH 
Fig. 7(a). Examples of some niosome forming surfactants: Compound 14,2 Compound 15,6 
Compound 16,9 and Compound 17.14 
18 
19 
Fig. 7(b). Niosomal membrane additives, Compound 18 = cholesterol, Compound 19 
Solulan C24.4 
110 Uchegbu & Schatzlein 
formation by non-ionic surfactants may be found in published reviews.4'63 Compounds 
such as Compounds 15 (from the sorbitan surfactant class) are established 
pharmaceutical excipients,64 and hence formulation scientists looking to prepare a 
niosome formulation for speedy transition to the clinic will do well looking at this 
class of molecules for exploitable materials. Most niosomes will not only contain the 
non-ionic surfactant, but will also contain other molecules such as the membrane 
stabilizer cholesterol [Fig. 7(b)].4 
The bilayer membrane is an ordered structure which may exist in the gel or 
liquid crystal state. Essentially, molecules are more mobile in the liquid crystalline 
state, enjoying lateral diffusion within the bilayer that is denied them in the gel state. 
For any system, the liquid crystal state exists at a higher temperature (T) than the 
gel state. An increase in temperature favors the transition from the gel to the liquid 
state because of the entropy gain (AS) associated with this transition, ultimately 
leading to a lowering of the free energy (AG) of the system. Cholesterol abolishes 
this membrane phase transition, thus fluidizing the gel state.65 
Niosomes are 30 nm to 120 JJLVSX in size4 and often their surfaces must be stabilized 
against aggregation. Molecules such as the cholesteryl poly(oxyethylene 
ether) — Solulan C246 (Compound 19, Fig. 7b) or the ionic molecule dicetyl 
phosphate66 have been used to confer steric and electrostatic stabilization on these 
vesicles respectively. The reader should be aware that the inclusion of minor 
quantities (<10% by actual weight or molar content) of ionic surfactant does not 
prevent these structures from being discussed in this chapter under the niosome 
heading. Niosomes are often formulated with minor quantities of cationic and other 
surfactants.4 
It can be said that the formulation of liposomes with poly(ethylene oxide) 
amphiphiles such as distearolyphosphatidylethanolamine-poly(ethylene glycol)26 
was the crucial step that allowed liposomes to become clinically relevant drug 
delivery systems. The resulting liposomes possess a hydrophilic polymer surface, 
which prevents recognition and clearance of the particles from the blood by the liver 
and spleen macrophages,26,67 thus increasing the liposomes' circulation time and 
allowing tumor targeting.68 Niosomes (non-ionic surfactant vesicles), when formulated 
with a water soluble poly(oxyethylene) cholesteryl ether — (Solulan C24), 
also circulate for prolonged periods in the blood, accumulate in the tumor tissue 
and improve tumoricidal activity.6 As well as stabilizing vesicles in the blood, 
poly(oxyethylene) amphiphiles also stabilize vesicles against aggregation, thus 
promoting vesicle colloidal stability.56 
Poly(oxyethylene) amphiphiles, such as Solulan C24, have a large hydrophilic 
head group [Fig. 7(b)], and are thus more hydrophilic than the vesicle forming 
amphiphiles, and hence the level of the former must be kept low to avoid solubilization 
of the membrane and the formation of mixed micelles.57 In actual fact, 
Vesicles Prepared from Synthetic Amphiphiles 111 
unusual morphologies57 result from the incorporation of non-micellizing quantities 
of Solulan C24 in vesicles as discussed below. 
4.2. Polyhedral vesicles and giant vesicles (Discomes) 
A series of unusual morphologies have been isolated from the hexadecyl diglycerol 
ether, Solulan C24, cholesterol phase diagram [Fig. 8(a)]. The addition of Solulan 
C24 to hexadecyl diglycerol ether [Compound 16, Fig. 7(a)] niosomes eventually 
results in the formation of mixed micelles.57 At sub-micellar concentrations of Solulan 
C24 (20-40 mole%), however, giant vesicles (discomes) of 25-100 pm in size are 
formed.57 Discomes are thermoresponsive vesicles, which become more leaky as 
the temperature is increased from room temperature to 37°C. These vesicles may 
thus be used to construct thermoresponsive controlled release systems. 
In cholesterol low regions of the hexadecyl diglycerol ether, cholesterol, Solulan 
C24 phase diagram, polyhedral vesicles [Figs. 8(a) and 8(b)] are found.9 These 
polyhedral vesicles are able to entrap water soluble solutes and the membrane, 
which is in the gel state contains areas of high and low curvature as shown in 
© Polyhedral Vesicles (2 -10 jim) 
® Spherical, helical, tubular Vesicles 
(0.5 -10 urn) 
•3 Discomes (10 - 30 jim) + small 
spherical & helical vesicles (0.5 -10 
jim) 
@ Discomes (12-60 fun) + mixed 
micelles 
 
\ Reverse Micelles 
••• \ 
\ 
V \ 
Oil 
Fig. 1. A hypothetical pseudo-ternary phase diagram of an oil/surfactant/water system 
with emphasis on microemulsion and emulsion phases. Within the phase diagram, existence 
fields are shown where micelles, reverse micelles or water-in-oil (w/o) microemulsions and 
oil-in-water microemulsions are formed along with the bicontinuous microemulsions. At 
very low surfactant concentrations two phase systems are observed (taken from Ref. 107). 
Recent Advances in Microemulsions as Drug Delivery Vehicles 129 
fractions, microemulsions are generally considered to be a dispersion of either oil 
or water droplets stabilized by an interfacial film of surfactant and where appropriate, 
cosurf actant. These droplet structures are probably the most commonly encountered 
type of microemulsion microstructure. It is worth noting that both an emulsion 
and a nanoemulsion can only occur in the form of a droplet, either as an oil-in-water 
or water-in-oil droplet. 
At intermediate oil and water compositions, it is obviously not possible for the 
microstructure to be composed of droplets of one phase dispersed in the other. 
In these cases, it is thought that a bicontinuous structure exists, in which the 
water and oil domains are separated by a regular or topologically chaotic continuous 
amphiphile-rich interfacial layer. A bicontinuous microemulsion is often the 
intermediate microstructure between an oil-in-water and a water-in-oil microemulsion, 
although a number of other microstructures such as cylinders and worm-like 
microemulsions have been reported to exist. 
In terms of its microstructure, a microemulsion is therefore a very complex 
system, and in instances where a microemulsion exists over a wide range of compositions, 
several different types of microstructure may be present.73 It is also important 
to remember that whatever the microstructure, a microemulsion is a dynamic 
system in which the interface is continuously and spontaneously fluctuating.104 
For this reason, microemulsions stabilized by polymeric surfactants may be the 
most long lived. 
1.4. Microemulsions, swollen micelles, micelles 
There is much debate in the literature as to what exactly differentiates a microemulsion 
from a micelle at low volume fractions of disperse phase. Some investigators 
have perceived a difference between microemulsions and micellar systems containing 
solubilized oil or water, and have used the terms "swollen" micellar solutions or 
solubilized micellar solutions to describe such systems. These investigators argue 
that the term microemulsion should be restricted to systems in which the droplets 
are of large enough size such that the physical properties of the dispersed oil or 
water phase are indistinguishable from those of the corresponding oil or water 
phase, thereby theoretically making it possible to distinguish between oil-in-water 
(or water-in-oil) microemulsions and micellar solutions containing small amounts 
of solubilized oil (water). However, in most cases, the transformation between 
micelles progressively swollen with oil (water) and a microemulsion containing 
an isotropic core of oil (water) appears to be gradual with no obvious transition 
point. As a consequence, there is no simple method available for determining the 
oil (water) content at which the core of the swollen micelle becomes identical to that 
of a bulk phase. Many researchers therefore use the term microemulsion to include 
130 Lawrence & Warisnoicharoen 
swollen micelles, but not micelles containing no oil (or water).34'107 In biotechnological 
applications, water-in-oil microemulsions are frequently known as reverse 
micelles and or even as nonaqueous media. 
1.5. Microemulsions and cosolvent systems 
The above broad definition does not require a microemulsion to contain any 
microstructure. In other words, it includes systems that are co-solvents, i.e. systems 
in which the constituent components are molecularly dispersed. Most researchers 
in the field agree, however, that for a microemulsion to be formed, it is important 
that the system contains some definite microstructure. In other words, there is a 
definite boundary between the oil and water phases, and at which the amphiphilic 
molecules are located and that a co-solvent is not a type of microemulsion. The only 
way to distinguish a microemulsion from a co-solvent unambiguously is to perform 
either a scattering study (light, X-rays or neutrons) or PFG-NMR measurements. 
Regions of co-solvent formation generally appear at low concentrations of oil or 
water. 
2. Microemulsions as Drug Delivery Systems 
It is clear from its description that microemulsions possess a number of properties 
that make their use as drug delivery vehicles particularly attractive. Indeed, 
microemulsions were first studied with the view of using them as potential vehicles 
for poorly-water soluble drugs, in the mid 1970s by Elworthy and Attwood.17 
However, it was not until the mid to late 1980s that they were widely investigated 
as drug delivery systems; this interest being largely the result of the arrival on the 
market of the cyclosporin A microemulsion preconcentrate, Neoral. 
Among the physical properties that make microemulsions attractive as drug 
delivery vehicle is their transparent nature, which means that the product is not 
only aesthetically pleasing, but allows easy visualization of any contamination. 
The small size of the domains present means that a microemulsion can be sterilized 
by terminal filtration.84 Furthermore, depending on the composition of the 
microemulsion, it may be possible to heat sterilize the microemulsions.39 Since oilin-
water microemulsions are able to incorporate lipophilic substances, they can 
be used to facilitate the administration of water-insoluble drugs.24 Significantly, 
the small droplet size provides a large interfacial area for rapid drug release, and 
so the drug should exhibit an enhanced bioavailability, enabling a reduction in 
dose, more consistent temporal profiles of drug absorption, and the protection of 
drug(s) from the hostile environment of the body. In addition to increasing the rate 
of drug release, microemulsions can also be used as a reservoir and actually slow 
the release of drug and prolong its effect, thereby avoiding high concentrations in 
Recent Advances in Microemulsions as Drug Delivery Vehicles 131 
the blood.64'142 Whether a drug is rapidly or slowly released from a microemulsion 
depends very much on the affinity of the drug for the microemulsion. Since 
microemulsions contain surfactants (cosurfactants) and other excipients, they may 
serve to increase the membrane penetration of drug.163'189 
A number of reviews have been presented, describing the pharmaceutical use of 
microemulsions.16'19-50'105"107'176 Since the last major review in the area was writen in 
2001, the present review will mainly deal with developments henceforth, although 
important work prior to this will be discussed when appropriate. 
2.1. Self-emulsifying drug delivery systems (SEDDS) 
Before discussing how microemulsions are being exploited in drug delivery, it 
is worth making one more distinction, namely the difference between a selfemulsifying 
drug delivery system (SEDDS) and a microemulsion. A SEDDS is a mixture 
of oil(s), and surf actant(s), ideally isotropic, sometimes containing cosolvent(s), 
which when introduced into aqueous phase under gentle agitation, spontaneously 
emulsifies to produce a fine oil-in-water dispersion.36'146 Typically, the size of the 
droplets produced by dilution of a SEDDS is in the range of 100 and 300 nm, while, 
upon dispersal in water, a SMEDDS formulation (a sub-group of the SEDDS) forms 
a transparent microemulsion with particle sizes <100 nm. ASMEEDS is also known 
as a pre-microemulsion concentrate.97 It is worth noticing that this method of producing 
a fine oil-in-water emulsion using a S(M)EEDS is identical to the low energy 
emulsification method for producing oil-in water nanoemulsions.173 It is therefore 
likely that a diluted S(M)EDDS and nanoemulsion are identically the same. 
The technique of low-energy or self-emulsification has been commercially 
exploited for many years in the agrochemical industry, in the form of emulsifiable 
concentrates of lipophilic herbicides and pesticides.146 However, it has only recently 
been introduced in the pharmaceutical industry as a tool to improve the delivery 
of lipophilic drugs by incorporating the drug into a S(M)EDDS formulation which 
is then filled into capsules.65 Once the capsule has been swallowed and its contents 
come into contact with the GI fluid, the drug containing (micro)emulsion should 
be spontaneously formed. Once the drug containing (micro)emulsion is formed, 
there should be little difference between the fate of the drug thus administered and 
the same drug administered in a (pre-formulated) microemulsion, although the 
droplets formed from the S(M)EDDS tend to be of a larger size. One advantage of 
administering a drug in a SMEEDS as opposed to a pre-formulated microemulsion, 
is its relatively small volume which can be incorporated into soft or hard gelatin 
capsules, convenient for oral delivery. 
To date, there has been a good amount of commercial success for the first selfmicroemulsifying 
drug delivery systems (SMEDDS) on the market, namely Neoral 
(cyclosporin A). In addition, the recent commercialization of two self-emulsifying 
132 Lawrence & Warisnoicharoen 
formulations, namely Norvir (ritonavir) and Fortovase (saquinavir), has undoubtedly 
increased the interest in SEDDS and other emulsion-based delivery systems 
to improve the delivery of a range of drugs of varying physico-chemical 
properties. 
However, there are a number of reasons why S(M)EDDS are not in greater 
widespread use, but the main reason is probably the stability of the diluted SEDDS, 
which is in fact a thermodynamically unstable emulsion (although it may exhibit 
some limited kinetic or "meta" stability). It should be noted however that as a 
SEDDS is either diluted just prior to administration or else in the body, the required 
droplet stability is less than 6 hrs (i.e. the transit time of materials down the small 
intestine). 
Although most studies of SEEDS have utilized isotropic liquids, the earliest 
reports of these self-emulsifying systems using pharmaceutical materials are in 
fact related to pastes based on waxy polyoxyethylene n-alkyl ethers.67 In the context 
of drug delivery via self-emulsifying systems, isotropic liquids are generally 
preferred to waxy pastes because if one or more excipient(s) crystallize(s) on cooling 
to form a waxy mixture, it is very difficult to determine the morphology of 
the materials. Despite this, there is currently a general move towards formulating 
semi-solid SEDDS. For example, attempts have been made to transform SEDDS into 
solid dosage forms by addition of large amounts of solidifying excipients such as 
adsorbents and polymers15,134 Unfortunately, as the ratio of SEDDS to solidifying 
excipients required for this approach is very high, this leads to problems in formulating 
drugs having limited solubility in the oil phase. Recent attempts have been 
made to reduce the amount of solidifying excipients by gelling the SEDDS with 
colloidal silicon dioxide.141 
Khoo et al.93 have recently reported the preparation of a halofantrine-containing 
lipid-based solid self-emulsifying system using either Vitamin E TPGS or a blend of 
Gelucire 44:14:Vitamin E TPGS as the base. Upon dispersal, these systems produced 
dispersions that the authors described as microemulsions. Studies in fasted dogs 
showed that these solid dispersions exhibited a five- to seven-fold improvement 
in absolute oral bioavailability, when compared with the commercially available 
tablet formulation. 
In a different approach, Nazzal et alP2 have determined the potential of a 
reversibly induced re-crystallized semi-solid self-nanoemulsifying drug delivery 
system, based on a eutectic interaction between the drug and the carrying agent, as 
an alternative to a conventional SEDDS. In these eutectic-based self-nanoemulsified 
systems, the melting point depression method allows the oil phase containing the 
drug itself to melt at body temperature from its semisolid consistency, and disperse 
to form emulsion droplets in the nanometer size range. Emulsion systems based on 
a eutectic mixture of lidocaine-prilocaine,135 and lidocaine-menthol87 have been 
Recent Advances in Microemulsions as Drug Delivery Vehicles 133 
used in the preparation of topical formulations. However, little is known of the use 
of eutectic mixtures for the preparation of self-(micro) emulsified formulations. 
2.2. Related systems 
There are a number of other putative delivery systems that are closely related to, 
or are prepared from, a microemulsion. These systems include a variety of gel 
formulations (including microemulsion-based gels, ringing gels, microemulsion 
gels) and double microemulsions. 
2.2.1. Microemulsion gels 
Oil-in-water microemulsions can be readily gelled or thickened by the addition of 
a non-interacting, water-soluble polymer such as polyHEMA,158 Carbopol 94044 or 
carrageenan179 to form clear "microemulsion gels". In these cases, it is the external 
aqueous phase that is gelled, while the microemulsion droplets are unperturbed. 
The structure of the resulting "microemulsion gel" is quite different, if it is prepared 
using an interacting polymer, such as stearate-polyethylene oxide-stearate. 
In this instance, the hydrophilic mid-block of the polymer is located in the continuous 
aqueous phase, while the hydrophobic end blocks are dissolved in the oil 
droplets, thereby connecting the various microemulsion droplets and resulting in 
the formation of a transient gel network.159 Clear, "microemulsion gels" are also 
sometimes obtained at surfactant and/or oil concentrations just outside the oilin-
water microemulsion region.180 Sometimes, the resultant gel "rings" or vibrates 
when tapped.180 The ringing is due to the resonance of shear modes within the gel 
body.167 Neither of these "microemulsion gels", which are water continuous, are 
true microemulsions, which are fluid by definition. 
Clear gels can also be formed in oil continuous systems. For example, a gel can 
be formed when water is added to reverse micellar solutions of lecithin-in-oil.5,12'161 
Here, the water causes the worm-like lecithin reverse micelles to intertwine and 
form a gel. In addition, gels, widely known as microemulsion-based gels, can be 
formed from water-in-oil microemulsions stabilized (predominately) by the dichain 
surfactant sodium bis (2-ethylhexyl) sulfosuccinate (AOT), when gelatin, the 
natural amphiphilic polymer is added.70,148 Microemulsion-based gels have now 
been prepared in systems in which a large amount of the AOT has been replaced 
by nonionic surfactant;88'89 or more recently, using in place of AOT, the single chain 
surfactant, cetyltrimethylammonium bromide in combination with pentanol.116 In 
these gels, the gelatin is thought to form water-continuous channels between the 
microemulsion-droplets. These microemulsion-based gels are very unusual in that, 
although they are oil continuous, they are electrically conducting. In addition, the 
134 Lawrence & Warisnoicharoen 
continuous oil behaves as if it were still a fluid, even though placing a gel in a 
solution of the oil does not dissolve it. 
All of these "microemulsion gels" have potential, or are being explored for use 
as drug delivery systems. Of particular interest is the fact that the gels possess the 
properties of being transparent, infinitely stable and readily prepared using only 
the mildest of mixing. In addition the wide range of microemulsion gels available 
means that it is possible to select the gel of the required consistency for application 
to large areas of skin, the nasal membrane, vaginal and buccal membranes and 
for permeation enhancement. Microemulsion-based gels have been explored as 
vehicles for the iontophorectic delivery of drugs.88 
2.2.2. Double or multiple microemulsions 
Double (or multiple) emulsions have attracted much interest as potential drug delivery 
vehicles. For example, adding a water-soluble drug into the internal aqueous 
phase of a water-in-oil-in-water emulsion may allow the sustained release of the 
water-soluble drug.59 A double microemulsion should offer similar advantages over 
the rate of drug release of entrapped solutes. Double emulsions are notoriously difficult 
to formulate due to the requirement to have one surfactant (or mixture of surfactants) 
to stabilize the first (internal) emulsion and a second surfactant (or mixture 
of surfactants) of quite different physico-chemical properties to stabilize the second 
emulsion. Although a few papers have detailed the production of nanoparticles 
from systems they described as double microemulsions,35,56'184 the term "double 
microemulsion" in this context is very misleading, as it refers to the mixing of two 
water-in-oil microemulsions of comparable composition, but containing different 
solute in the aqueous phase. 
There are however two papers which describe the preparation of double (oilin-
water-in-oil) microemulsions. In the first, Castro et al.,30 report spectroscopic 
studies of nifedipine in Brij 96 based oil/water/oil multiple microemulsions. In the 
second, Carli et al.,29 detail the preparation of an oil-in-water-in-oil microemulsion 
from an oily phase of either polyglycolized glycerides or a mixture of mono-, diand 
tri-glycerides, which is microemulsified using a mixture of water and surfactant 
(soy lecithin and Tween 80). The resultant o /w microemulsion is subsequently redispersed 
in an oily phase to produce the double (o/w/o) microemulsion.29 
2.3. Processed microemulsion formula tions 
2.3.1. Solid state or dry emulsions 
In practical terms, a solid dosage form is preferable to a liquid dosage form in respect 
of convenience, ease of handling and accurate dosing. Consequently, a number of 
Recent Advances in Microemulsions as Drug Delivery Vehicles 135 
researchers have attempted to develop powdered, re-dispersible emulsion-derived 
formulations, known as solid state or dry emulsions. Such solid-state emulsions 
can be used to modulate the release rates of emulsified compound.128 Dry emulsions 
have been variously prepared by removing water from an oil-in-water emulsion, 
using water-soluble182 or -insoluble150 solid carriers or indeed a mixture 
of both water-soluble and water-insoluble carriers,143 by rotary evaporation,128 
lyophilization,115 or spray drying.55'127 
Attempts have also been recently made to prepare dry microemulsions using 
similar methodology. For example, Moreno et al.126 have lyophilized an amphotericin 
B-containing lecithin-based oil-in-water microemulsion in the presence 
of 5wt% mannitol. The lyophilized product was an oily cake from which the 
microemulsion could be easily reconstituted over several months. The rationale 
for developing the water-free formulation was to avoid the hydrolysis of lecithin, 
which occurs upon its dispersal in water, thereby preventing any deterioration of 
the formulation upon storage. Overall, the lyophilized lecithin based oil-in-water 
microemulsions appear to be valuable systems for the delivery of amphotericin B, 
with regard to ease and low-cost of manufacturing and their stability and safety, 
compared with other formulations already in the market. 
In a recent paper, Carli et al.29 reported an alternative approach to prepare a 
"dry" formulation known as a nanoemulsified composite of the coenzyme Q10, 
ubidecarenone. This composite is prepared by incorporating the ubidecarenone 
into the inner phase of the double microemulsion, which is then deposited onto a 
solid microporous carrier such as cross-linked polyvinylpyrrolidone. Among the 
advantages offered by this approach are good processing and storage properties, 
easy re-dispersibility in water, high bioavailability and maintenance of the submicron 
size of the released droplets. 
Kim et al.,97 have prepared entric-coated solid state premicroemulsion concentrates 
by first preparing a pre-microemulsion concentrate containing 10 wt% of the 
drug cyclosporin A, 18.5 wt% of a medium chain triglyceride, 51 wt% of surfactant 
and 20 wt% of cosurfactant. The pre-microemulsion concentrates were then 
enteric-coated as films using polymers, such as sodium alginate, Eudragit L 100 
and cellulose acetate phthalate, and the resulting films were pulverized to produce 
powdered, dry, enteric coated premicroemulsion concentrates. Using this approach, 
the authors successfully prepared a once-a-day oral dose form of cyclosporin A. 
3. Formulation 
Microemulsions are far more difficult to formulate than emulsions because their formation 
is a highly specific process involving spontaneous interactions among their 
constituent molecules. In addition, in a number of cases, effects due to the order of 
136 Lawrence & Warisnoicharoen 
the mixing of the component molecules have been observed. Since no adequate theory 
currently exists to predict from which molecules microemulsions can be formed, 
mainly because of the requirement to determine a number of unknown parameters, 
microemulsion formulations are generally developed empirically, although some 
useful practical guidance as to the choice of the constituent components can be 
found in the literature.105,107 
A recognized and classical approach to microemulsion formulation is to undertake 
a systematic study of the phase behavior of the systems understudying utilizing 
of phase diagrams. A major drawback of this approach is the considerable time 
it takes to develop the phase diagram, especially considering the combination of 
possible oil, surfactant and cosurfactant, and the fact that time may be necessary for 
a system to equilibrate. Heat and sonication are therefore often used, particularly 
with systems containing nonionic surfactants, to speed up the formation process. 
While there are now commercially available automated systems to prepare phase 
diagrams,78 the chief drawback of these systems is their cost. 
A number of attempts have been recently made to use modeling to predict 
microemulsion formation, thereby aiding in the formulation of microemulsions. 
A range of modeling techniques have been used including artificial 
neural networks,3'4'8'149 genetic algorithms3,174 and a combination of data mining, 
computer-aided molecular modeling, descriptor calculation and multiple linear 
regression techniques.174,175 Unfortunately, however, all of these techniques 
require a considerable amount of work prior to prediction, thereby restricting 
their potential usefulness. Furthermore, the amount of work required for the predictions 
increases as the number of components of the microemulsion increase; 
microemulsions formulated from five components (i.e. oil, water, surfactant, cosurfactant, 
electrolyte and drug) are not uncommon in pharmaceutical use. To the 
authors' knowledge, to date, no work has been performed predicting how much 
drug can be incorporated into a microemulsion and whether the presence of drug 
has any effect on microemulsion phase behavior. This is an important ommision 
as microemulsions cannot be considered to be inert, since the presence 
of drug in some instances (greatly) influences phase behavior (see for example 
Ref. 138). 
3.1. Surfactants and cosurfactants 
The selection of components for the preparation of microemulsions suitable for 
pharmaceutical use involves a consideration of their toxicity, and if the systems 
are to be used topically, their irritancy and sensitizing properties as well. There are 
a number of surfactant and cosurfactants that are considered acceptable for use 
as excipients in pharmaceutical formulation.153 Strickley172 has recently reviewed 
Recent Advances in Microemulsions as Drug Delivery Vehicles 137 
those surfactants and cosurfactants currently used in commercially available oral 
and intravenous formulations. 
In the general scientific literature, by far the most widely used surfactant to prepare 
a microemulsion is the double chain, ionic surfactant, sodium bis (2-ethylhexyl) 
sulfosuccinate (AOT), although a large number of studies have used the single 
chain, nonionic surfactants of the type QEj, where i is the number of carbons in the 
alkyl chain, C, and j is the number of ethylene oxide units in the polyoxyethylene 
chain, E. Both AOT and the QEj surfactants possess the important advantage of 
being able to form microemulsions in the absence of a cosurfactant,121'185 unlike 
most other types of surfactant such as the widely studied single chain, ionic surfactant, 
sodium dodecyl sulphate (SDS), which will only form microemulsions in the 
presence of an alcohol cosurfactant. Neither AOT nor SDS would be considered to 
be apprpropriate for the preparation of pharmaceutically acceptable microemulsions, 
even though they are listed in the Pharmaceutical Excipient Handbook, 
Rowe et al.l5i 
As a general rule, nonionic and zwitterionic surfactants tend to be less toxic 
than ionic surfactants and are therefore more widely used as pharmaceutical 
excipients.154 Assuming that the surfactants do not degrade into toxic materials, 
surfactants that posses biodegradable/chemically unstable linkers tend to exhibit 
less chronic toxicity than those that are chemically stable. For example, as a group, 
the polyoxyethylene n-acyl surfactants exhibit ~ten times less chronic toxicity than 
their n-alkyl counterparts, mainly due to their quicker degradation time of days as 
opposed to weeks. When it comes to comparing acute toxicity, the two groups of 
surfactants exhibit comparable toxicity. 
Perhaps the most widely used nonionic surfactants in pharmaceutical formulations 
are the polyoxyethylene sorbitan n-acyl esters, i.e. the Tweens and in particular, 
Tween 20 and Tween 80, both of which are used parenterally and orally. In addition, 
polyoxyethylene derivatives of the triglyceride, castor oil have acceptability 
for intravenous administration. Other pharmaceutically acceptable surfactants are 
the polyoxyethylene n-alkyl ethers and n-acyl esters, although both these groups of 
surfactant tend to be restricted for topical use.154 Other nonionic surfactants that are 
currently attracting much pharmaceutical interest, although they do not yet have 
acceptability, are the polyglycerol n-acyl esters, the n-alkyl amine N-oxides and 
the w-alkyl polyglycosides (or sugar surfactants). The n-alkyl polyglycosides have 
attracted much pharmaceutical interest, not because of their excellent biodegradability, 
but because they can be manufactured from renewable resources. All of the 
aforementioned surfactants have been used to prepare microemulsions, generally 
as sole surfactant, the only exception being the w-alkyl polyglycosides, which tend 
to require the presence of a cosurfactant. 
Pluronics (or poloxamers) of the type poly(ethylene oxide)-poly(propylene 
oxide)-poly(ethylene oxide) (PEO-PPO-PEO) are another class of pharmaceutically 
138 Lawrence & Warisnoicharoen 
attractive surfactant. Interestingly, most reports detailing the use of polymeric surfactants 
to stabilize a microemulsion describe the preparation of water-in-oil, generally 
in conjunction with a second surfactant.96,181 Siebenbrodt and Keipert162 have 
reported the formation of a triacetin-in-water microemulsion using Pluronic L 64 
as sole surfactant. Lettow et al.,m have used Pluronic PI23 as sole surfactant to 
prepare oil-in-water microemulsions, incorporating a 1:1 oil:P123 weight ratio of 
either 1,3,5-trimethylbenzene or 1,2-dichlorobenzene. 
Finally, the pharmaceutically acceptable zwitterionic lecithin has been extensively 
used as a surfactant, however, with very few exceptions, it is not possible to 
prepare a microemulsion using lecithin as sole surfactant. Generally, lecithin is combined 
with another surfactant such as Tween 80, or a cosurfactant such as ethanol, 
when formulating microemulsions. 
Although ethanol is considered to be pharmaceutically acceptable, typical 
cosurf actants such as propanol and butanol are not. In addition to toxicity issues, the 
use of such cosurfactants, which may possess partial oil and water-solubility, can 
lead to problems with the dilutability of the microemulsion. This is a particular issue 
if the microemulsion is to be administered orally or parenterally. Consequently, a 
number of researchers have explored the use of a second surfactant as cosurfactant 
when formulating a microemulsion. Microemulsions thus prepared tend to be very 
stable against dilution, as the "cosurfactant" generally has little solubility in either 
the oil or aqueous phase. Alternately (pharmaceutically acceptable), short chain 
mono- and di-glycerides have been used in place of a short chain alcohol to successfully 
prepare microemulsions. In a number of instances, short chain fatty acids such 
as sodium caprylate have been used as cosurfactants, primarily for the formation 
of microemulsions for oral delivery; sodium caprylate is known to enhance absorption 
of drugs across the gastrointestinal tract. A number of researchers have also 
used cosolvents such as the polyhydric alcohols, sorbitol, glycerol and propylene 
glycol to aid microemulsion formation. In a number of instances, these materials 
have been described as "cosurfactants", which quite clearly do not sit in the interfacial 
surfactant monolayer. Rather, they tend to exert their effect by altering the 
solvent properties of the polar phase. 
3.2. Oils 
Most reports in the chemical literature detail the preparation of microemulsions 
using aromatic oils such as benzene and short chain alkanes such as hexane. "Pharmaceutical" 
oils, unlike those used in the chemical and agrochemical industries, 
tend to be large in terms of molecular weight and therefore volume, and are relatively 
polar. Both of these properties tend to work against the oil when it comes to 
formulating it in a microemulsion, as it is well established that smaller molecular 
volume oils are easier to solubilize and are solubilized to a greater extent than larger 
Recent Advances in Microemulsions as Drug Delivery Vehicles 139 
oils.2 Although there are reports that in some systems, particularly those containing 
surfactants with long, unsaturated hydrophobes such as polyoxyethylene (10) oleyl 
ether, the largest molecular volume oil is solubilized to a greater extent than some of 
the smaller molecular volume oils.122 The most commonly used "pharmaceutical" 
oils are medium and long chain triglycerides, and esters of fatty acids such as ethyl 
oleate, isopropyl myristate are popular. 
It has become common practice for researchers to screen the solubility of drug 
in the various components of the microemulsion, in order to predict the optimal 
composition of the final formulation. However, extreme care has to be exercised 
when using this approach, as very often, the solubility in the final microemulsion 
formulation does not correlate well with that seen in the various components. 
3.3. Characterization 
It is noticeable that in contrast to their ease of preparation, it is very difficult to 
establish the microstructure of a microemulsion. Yet, such information is important 
as it may influence the drug behavior of the microemulsion in use. For example, it 
is known that the microstructure of the microemulsion may alter the release rate of 
any incorporated solute.1'95 
Currently, a range of physico-chemical techniques are used to characterize 
microemulsions. These techniques are often used in tandem to obtain a better picture 
of the system, as it is unlikely that any one technique alone will give sufficient 
information.144 Scattering techniques (light, neutron and X-ray) and pulsed 
field gradient NMR are generally used to determine the microstructure of the 
microemulsion. One serious limitation with characterizing microemulsions is that 
most techniques rely on the concentration of disperse phase being low enough to 
avoid particle-particle interactions, as an estimated volume fraction of 1 vol% is 
suitable.123 The requirement is a particular problem with microemulsions that contain 
cosurfactants that partition between the oil and water phases, because these 
systems frequently undergo a change upon dilution. 
4. Routes of Administration 
Although most of the original work exploring microemulsions as drug delivery 
vehicles examined their potential for oral drug delivery, microemulsions have now 
been explored as vehicles for most routes of administration. Currently, they are 
probably most widely studied for their potential as transdermal delivery vehicles. 
4.1. Oral 
Microemulsions (and SMEDDS) have been widely studied as oral drug delivery 
vehicles. Indeed, the first commercially available "microemulsion" formulation was 
140 Lawrence & Warisnoicharoen 
a premicroemulsion concentrate of the lipophilic peptide, cyclosporin A. This formulation, 
known commercially as Neoral, was introduced onto the market in the 
late 1980s and immediately attracted much attention, mainly because of the high 
and reproducible bioavailability it produced, but also because developments in 
biotechnology at that time meant that it had never been easier to produce on a large 
scale therapeutically-relevant protein and peptides. Unfortunately, because of their 
physico-chemical properties, in particular their large size and poor stability, proteins 
and peptides are very difficult to formulate. Microemulsions offered an attractive 
solution to this problem, and consequently, most of the original exploratory 
studies on microemulsions as drug delivery vehicles were spent developing oral 
protein/peptide microemulsion formulations. 
4.1.1. Proteins and peptides 
As the majority of therapeutic proteins and peptides are hydrophilic and watersoluble, 
most studies utilizing microemulsions as vehicles for such molecules have 
exploited water-in-oil microemulsions. After cyclosporin A, which is unusually 
highly lipophilic, for a therapeutic peptide, the most widely studied peptide is 
insulin, with much of the early work in this area being performed by Ritschel.152 
For example, Kraeling and Ritschel101 compared the peroral microemulsion formulation 
of insulin and capsule forms and determined that the microemulsion 
formulation increased the bioavailability of the insulin. Recently, more complex 
microemulsion-based systems have been developed in an attempt to improve 
the extent of insulin absorption. For example, a recent study performed by 
Natnasirichaiku et al.186 showed a significant improvement in the oral bioavailability 
of insulin (in diabetic rats) when administered in nanocapsules dispersed 
in a water-in-oil microemulsion. Santiago et al.155 have developed a new, enteric 
oral dosage form of insulin, in which an association of insulin and cyclodextrin 
contained within a microemulsion is processed into granules. In the most recent 
study aimed at developing an oral formulation of insulin, Iek et al.77 used a conventional 
lecithin-based water-in-oil microemulsion formulation prepared from 
21.6 wt% water, 37.6 wt% Labrafil M 1944 CS as oil and stabilized by 40.8 wt% of a 
1:1 weight ratio of lecithin (Phospholipon 90G) and ethanol. In addition to insulin 
(21.6IU/g water), some of the microemulsions contained the enzyme inhibitor 
aprotinin (2500KlU/g water). Although it is the first time that a microemulsion 
formulation has contained both a protein/peptide and an enzyme inhibitor, the 
concept of adding an enzyme inhibitor, to a formulation containing a peptide in an 
attempt to reduce its degradation is not new.188 The plasma glucose and insulin levels 
of the rats after intragastric administration of the formulations to both diabetic 
and non-diabetic rats were significantly different from those obtained after oral 
Recent Advances in Microemulsions as Drug Delivery Vehicles 141 
administration of an aqueous insulin solution. Although the addition of aprotinin 
to the microemulsion containing insulin increased bioavailability when compared 
with those not containing it, the difference was insignificant. 
Other peptides formulated as water-in-oil microemulsions in an attempt to 
improve their oral absorption include RGB peptides,37'38 and more recently, Nacetylglucosaminyl-
N-acetylmuramyl dipeptide (GMDP).119 The poor bioavailability 
of GMDP has been attributed to both its poor stability in the lumen of the 
gastrointestinal tract and its poor intestinal permeability. When GMDP was administered 
intraduodenally in a water-in-medium-chain trigylceride microemulsion, a 
ten-fold increase in bioavailability was observed, i.e. a bioavailability of 80.2% was 
achieved as opposed to 8.4%, seen after administration of an aqueous solution of 
GMDP. This increase in bioavailability is consistent with the work of Constantinides 
et al.37,3S who utilized a similar medium chain triglyceride based microemulsion to 
increase the oral bioavailability of the water-soluble peptide SK&F 106760, after 
intraduodenal administration to rats. 
Ke et al.92 have recently reported an attempt to develop water-in-oil microemulsions 
suitable for the incorporation of therapeutic proteins and peptides using 
a medium chain triglyceride, water and tocopheryl polyethylene glycol 1000 
succinate (TPGS) as the primary surfactant. However, as TPGS could not form 
microemulsions when used as sole surfactant, it was mixed with a second surfactant, 
either Tween 20,40,60 or 80, at a weight ratio in the range of 4:1 to 1:4. A range 
of glycols and polyols were examined as cosurfactants. Although stable, transparent 
microemulsion and gel regions were identified, the extent of these regions was 
influenced by the precise nature and the amount of the secondary surfactant and 
cosurfactant. For example, Tween 80, which is an ester of the unsaturated CI 8 fatty 
acid, oleic acid, was more effective in forming a microemulsion than Tween 60, 
which is an ester of the saturated C18 fatty acid, stearic acid. In this study, although 
the microemulsions were ultimately intended for use as delivery vehicles for protein 
or peptide drugs, they were not examined for this purpose. 
4.1.2. Other hydrophilic molecules 
Other water-soluble therapeutic molecules that have been administered in 
microemulsions include the aminoglycoside antibiotic, gentamicin74 and the biologically 
active polysaccharide, heparin." In common with all aminoglycosides, 
gentamicin is highly polar and is therefore considered unlikely to be absorbed 
from the gastrointestinal tract via simple diffusion. In order to facilitate the transmucosal 
delivery of the drug, Hu et al74 prepared a SMEDDS formulation of gentamicin 
using a range of surfactants. When Labrasol was used as surfactant, a 54.2% 
bioavailability of gentamicin was obtained, compared with only 8.4 and 3.4% when 
142 Lawrence & Warisnoicharoen 
Tween 80 and Transcutol P were respectively used. Labrasol was also found to 
inhibit intestinal secretory transport from the intestinal enterocytes, providing the 
formulation with the additional benefit of inhibiting the efflux of gentamicin out of 
the enterocytes into the GI lumen. 
Due to its low bioavailability, heparin is generally administered by injection. In 
an attempt to formulate an orally active version of heparin, Kim et al." synthesized 
a low molecular weight heparin (LMWH)-deoxycholic acid (DOCA) conjugate 
(termed LMWH-DOCA) and formulated it in a water-in-oil microemulsion using 
as oil, the medium chain trigylceride, tricaprylin, a mixture of Tween 80 and Span 
20 surfactants, LMWH-DOCA and water (volume ratios of 5:3:1:1 respectively). 
Oral administration of LMWH-DOCA in the water-in-tricaprylin microemulsion 
to mice resulted in a bioavailability of 1.5%. Toxicity studies suggested that the 
enhancement in bioavailability, observed with the DOCA-conjugated LMWH, was 
administered in a microemulsion not due any local toxicity such as disruption or 
damaging of the intestinal membrane. 
4.1.3. Hydrophobic drugs 
A number of poorly water-soluble, low molecular weight, lipophilic drugs have 
also been formulated in microemulsions (or SMEEDS) for oral delivery including 
nitrendipine,90 danzol145 halofantrine94 and biphenyl dimethyl dicarboxylate.98 
These studies serve as an illustration of how important it is to understand the 
influence on microemulsion formation of the various formulation components. It 
is worth commenting that the main use of SMEEDS formulations is for the oral 
administration of lipophilic drugs. 
Formulating nitrendipine in a SMEEDS formulation, composed of a 1:1 (w/w) 
mixture of glycerol monocaprylic ester (MCG) and propyleneglycol dicaprylic ester 
(DCPG) and nonionic surfactant (various), was observed to significantly enhance 
its absorption when compared with a suspension or an oil solution,90,91 and served 
to reduce the effect of the presence of food on its absorption. However, the absorption 
profile of nitrendipine was seen to vary with the type of surfactant used; 
absorption was rapid from the Tween 80-stabilized formulation, while the HCO-60- 
based formulation gave a prolonged plasma concentration profile. No absorption of 
nitrendipine was observed from the formulation containing BL-9EX (polyoxyethylene 
alkyl ether, C12E9). Damage to the gastrointestinal mucosa also differed with 
the type of surfactant employed. HCO-60 and Tween 80-based formulations were 
mild to the organs, while BL-9EX-based formulation caused serious damage. 
The study of Porter et al.145 appropriately demonstrates the effect of changing 
the nature of the trigylceride involved in the formulation on drug absorption. 
These workers studied three lipid-based danazol formulations; namely a long-chain 
triglyceride solution (LCT-solution), a SMEDDS based on long (C18) chain lipids 
Recent Advances in Microemulsions as Drug Delivery Vehicles 143 
(LC-SMEDDS) and a SMEEDS formulation containing medium (C8-C10) chain 
lipids (MC-SMEDDS). These formulations were administered to fasted beagle dogs 
and their absorption, compared with that obtained with a micronized danazol formulation 
administered postprandially and in the fasted state. Although both the 
LCT-solution and LC-SMEDDS formulations were found to significantly enhance 
the oral bioavailability of danazol, when compared with fasted administration of the 
micronized formulation, the MC-SMEDDS produced little improvement in danazol 
bioavailability. This result was partly attributed to the fact that upon digestion of 
the medium-chain formulation, significant drug precipitation was observed. 
Khoo et al.94 also considered the effect of formulating halofantrine as a 
pre-microemulsion concentrate in a formulation based on either a medium- or longchain 
triglyceride. Both formulations, which were administered as soft-gelatin capsules, 
contained the same amount of medium or long chain trigylceride and were 
stabilized by the same surfactant/cosurfactant mixture, consisting of Cremophor 
EL and ethanol. Although the plasma levels of the drug were not significantly 
different between the two formulations, the amount of drug absorbed lymphatically 
varied in that 28.3% of the dose administered in the long-chain trigylceride 
formulation was transported lymphatically, as opposed to only 5.0% of the dose 
administered in the medium-chain formulation. 
Kim et al.9S attempted to improve the solubility and bioavailability of biphenyl 
dimethyl dicarboxylate, a drug used in treating liver diseases, by formulating it as 
a premicroemulsion concentrate. In order to optimize drug loading in the formulation, 
these workers screened drug solubility in a range of surfactants and oils, 
and on the basis of these results selected: Tween 80 and Neobee M-5. However, care 
must be taken when using this approach to optimize the formulation with respect to 
drug loading, as it has shown that solubility of drug in the bulk components is not 
a reliable indicator of solubility, in the final microemulsion formulation.120,122 The 
danger of predicting drug solubility in the final formulation, on the basis of bulk 
solubility, can be seen in the study of Kim et al.98 where the solubility of the drug 
in a formulation consisting of a 2:1 weight ratio of Tween 80 to Neobee M-5 was 7 
times that of the formulation containing a Tween 80:Neobee M-5 weight ratio of 1:4, 
despite the solubility of the drug in Neobee M-5 being 10 times that seen in Tween 
80. The final formulation, which consisted of 35 wt% triacetin and 65 wt% Tween 
80 and Neobee M-5 at a weight ratio of 2:1, greatly enhanced the oral bioavailability 
of BDD, possibly due to the increased solubility of the drug and its immediate 
dispersion in the gastrointestinal tract. 
Itoh et al.79 optimized the formulation of the poorly water-soluble 
drug N-4472, N-[2-(3,5-di-tert-butyl-4-hydroxyphenethyl)-4,6-difluorophenyl]-N- 
[4-(Nbenzylpiperidyl)] urea, by complexing it with L-ascorbic acid and incorporating 
the complex into a SMEEDS comprising Gelucire 44/14, HCO-60 and sodium 
dodecyl sulfate. Upon dilution with water, the SMEEDS formulation produce a fine 
144 Lawrence & Warisnoicharoen 
dispersion of 18 nm droplets which were stable over the pH range of 2.0 to 7.0. The 
oral bioavailability of the drug was between 2-4 times that which was obtained 
with an aqueous solution of the complex. 
4.2. Buccal 
To date, very little work has been performed on investigating the use of microemulsions 
as vehicles for buccal delivery. In 1988, Ceschel et a\?x showed that the penetration 
of the essential oil, Salvia sclarea L. through porcine buccal mucosa in vitro 
was increased when formulated as a microemulsion, as opposed to the pure essential 
oil. Scherlund et al.5S investigated the potential of lidocaine and prilocaine 
thermosetting microemulsions and mixed micellar solutions as drug delivery systems 
for anesthesia of the periodonlal pocket. The formulations contained between 
2-10 wt% of a eutectic mixture of lidocaine or prilocaine (melting point 18°C), while 
the block copolymer surfactants, Pluronic F127 and F68, were present at between 
13 and 17 wt% for F127, and between 2 and 6 wt% for F68. F127 was chosen, as it is 
known to gel at body temperature and it is important that the formulation is easy 
to apply, remain at the application site, have a fast onset time, be non-irritant, and 
stable under normal storage conditions. The pH of the formulations was varied 
between 5 and 10. Most of the combinations were found to result in clear solutions, 
presumably oil-in-water microemulsions or mixed micellar solutions, depending 
on the pH of the system. At low pH, lidocaine and prilocaine are positively charged, 
and they could be expected to behave largely as water-soluble cationic surfactants, 
hence possibly forming mixed micelles. On the other hand, at high pH, the drug 
substances are poorly soluble and could be expected to act largely as hydrophobic 
solutes and form the core of the microemulsion droplets. 
4.3. Parenteral 
In recent years, considerable emphasis has been given to the development of 
injectable microemulsions (o/w) for the intravenous delivery of drug, in order 
to increase the solubility of the drug39'138'139 to reduce drug toxicity,25'26'126 to 
reduce hypersensitivity,72 and to improve drug solubility and reduce pain upon 
injection.109 A very recent development is the formulation of microemulsions as 
long circulating vehicles, and more recently, as drug tageting agents. In addition, 
water-in-oil microemulsions have been investigated as depot vehicles for the intramuscular 
delivery of drugs.22'64 
The first published study which established the potential of microemulsions 
for use in intravenous delivery was probably that of von Corswant 
et al. in Ref. 39. These researchers prepared a pharmaceutically acceptable, 
bicontinuous microemulsion from a medium-chain triglyceride oil, poly(ethylene 
Recent Advances in Microemulsions as Drug Delivery Vehicles 145 
glycol) 400 and ethanol cosolvents and stabilized by soybean phosphatidylcholine 
and poly(ethylene glycol)(660)-12-hydroxystearate. Prior to administration, the 
microemulsion required dilution with a suitable aqueous phase. Upon dilution, the 
microemulsion formed an oil-in-water microemulsion with droplets of size between 
60 and 200 nm, smaller than the size of the droplets in a commercial intravenous 
emulsion, namely Intralipid. From their animal studies, the authors concluded that 
the microemulsion they developed was suitable for administion by intravenous 
infusion to conscious rats. Unfortunately, although the researchers did determine 
drug solubility in the bicontinuous microemulsions, they did not report this. 
Park and Kim138 also investigated the formulation of poorly water-soluble 
flurbiprofen at ~8 times its aqueous solubility into an oil-in-water microemulsion 
suitable for intravenous administration. The microemulsions were prepared 
using varying weight ratios of oil (ethyl oleate) to surfactant (Tween 20), and contained 
a range of isotonic solutions as the polar (aqueous) phase. Unfortunately, 
insufficient information was supplied regarding the precise compositions of the 
microemulsions, in particular, how much oil and surfactant were present, so as to 
draw conclusions about the formulation; (perhaps surprisingly) the ratio of oil to 
surfactant used did not seem to have any effect on the amount of drug solubilized 
and that the presence of too much drug had a destabilizing effect on the microemulsion. 
Disappointingly, the pharmacokinetic parameters of flurbiprofen, after intravenous 
administration of flurbiprofen-loaded microemulsion to rats, were also 
not significantly different from those of flurbiprofen in phosphate buffered saline 
solution. In a later publication, Park et a/.138 overcame the problem of stability 
seen in their earlier study by replacing the surfactant Tween 20 with lecithin and 
distearoylphosphatidyl- ethanolamine-N-poly(ethyleneglycol) 2000 (DSPE-PEG) 
and using ethanol as a cosolvent. Due to the presence of the long chain polyoxyethylene 
groups on the exterior surface of the microemulsion droplets, it was 
perhaps unsurprising that the biodistribution of flurbiprofen administered in this 
microemulsion was quite different. In particular, reticuloendothelial uptake of flurbiprofen 
decreased, suggesting that it may ultimately be possible to target drugs 
incorporated in this microemulsion to different sites of the body. 
As part of a series of papers, Brime et al.25,26 and Moreno et al.126 prepared a 
novel amphotericin B lecithin-based oil-in-water microemulsion, in an attempt to 
produce a formulation with less toxic effects than the currently available commercial 
formulation, Fungizone. The microemulsion which contained as oil isopropyl 
mystriate and a mixture of either Tween 80 or Brij 96 with lecithin as surfactant. 
In some instances, formulation was lyophilized in an attempt to increase its stability. 
The overall results of the toxicity studies were encouraging as the amphotericin 
B-containing microemulsions exhibited a low toxicity, suggesting a potential 
therapeutic application. 
146 Lawrence & Warisnoicharoen 
Zhang et al.m prepared a lecithin-based SMEDDS formulation of the drug 
norcantharidin. Upon dilution, the release rate of norcantharidin contained in the 
SMEEDS formulation was found to be dependent on the size of the disperse phase 
and the type of lecithin used. Interestingly, although norcantharidin was poorly 
soluble in the ethyl oleate and only slightly soluble in water, microemulsions containing 
ethyl oleate oil exhibited a significant increase in solubilization over the 
corresponding aqueous solution. 
Clonixic acid is currently marketed in salt form because of its poor watersolubility. 
However, the commercial dosage form causes severe pain after intramuscular 
or intravenous injection. To improve the apparent aqueous solubility of 
clonixic acid and to reduce the pain it causes on injection, Lee et al. (2000) incorporated 
3 mg/mL clonixic acid into oil-in-water microemulsions (size 120 nm) prepared 
from pre-microemulsion concentrate of castor oil, and a mixture of Tween 
20 and Tween 85 surfactants (present in a weight ratio of 5:12:18). Although the 
microemulsion formulation significantly reduced the number of rats licking their 
paws as well as the total licking time, suggesting less pain induction by the 
microemulsion formulation; the pharmacokinetic parameters of clonixic acid after 
intravenous administration were not significantly different from those of the commercial 
formulation, lysine clonixinate. The results of the study suggested that a 
microemulsion formulation is an alternative vehicle for clonixic acid. 
Paclilaxel (Taxol) injection is known to cause hypersensitivity reactions. Consequently, 
He et al.72 explored whether it was possible to prepare a non-sensitizing 
paclitaxel microemulsion using egg phosphatidylcholine, Piyronic F68 ancl Cremophor 
EL as surfactants, and ethanol as cosurfactant. Note that there was no 
mention of the presence of a specific oil. The study showed that for an equivalent 
dose, the paclitaxel microemulsion did not cause any hypersensitivity reaction, 
whereas Taxol did. In addition, the bioavailability of the paclitaxel in the new 
microemulsion was significantly higher and the elimination rate slower than that 
achieved with Taxol. The authors suggested that the drug molecules, trapped in the 
oil droplets, diffused into the systemic circulation slowly. Furthermore, the small 
particle size of the droplets (10-50 nm) meant that the microemulsion droplets could 
escape from uptake and phagocytosis of RES. Infact it was previously suggested 
that it should be possible to modify the surface of the microemulsion droplets, with 
polyoxyethylene chains, to significantly improve circulation time.57'118'190 
Kanga et al.S6 have recently explored the possibility of optimizing the release 
of paclitaxel from a SEEDS formulation using the polymer, PLGA. The SEEDS formulation, 
which was a mixture of drug, tetraglycol, Cremophor ELP, and Labrafil 
1944 also contained PLGA of varying molecular weight. The droplet size of the 
microemulsions was in the range of 45-270 nm, with the systems without PLGA 
exhibiting the smaller size. The release rate of paclitaxel decreased in the order of 
Recent Advances in Microemulsions as Drug Delivery Vehicles 147 
PLGA, PLGA 8 K, PLGA 33 K, and PLGA 90 Kg/mol, suggesting that the molecular 
weight of PLGA in microemulsion could control the release rate of paclitaxel from 
microemulsion. 
4.3.1. Long circulating microemulsions 
Long circulating microemulsions have been suggested as an alternative formulation 
to long circulating vesicles on the basis of their small size, thus avoiding uptake by 
the RES, their stability and their ability to solubilize lipophilic compounds more 
effectively than vesicles, and their ease of preparation. 
Wang et al.,S3 and Junping et al.,m have determined the potential of intravenous 
delivery systems of emulsion/microemulsion systems based on vitamin E, cholesterol 
and PEG2ooo-lipid. In their first study, Wang et a/.,183 prepared emulsions containing 
1 part drug, 3 parts vitamin E, 3 parts cholesterol and 3 parts PEG2000-DSPE 
with the final formulation containing 5mg of drug in 10 mL of saline solution. 
Although the emulsion was reported to form spontaneously on the addition of the 
required amount of saline, the formulation was homogenized to produce a more 
uniform particle size distribution of 123.0 ±1.2 nm; no information was given as to 
the size of the droplets prior to homogenization. The zeta potential and drug loading 
efficiency of the sub-micron emulsion were -12.67 + 1.35 mv and 96.3 + 0.3. 
Although the size and loading efficiency of the formulation remained uncharged 
when stored at 7 to 8°C for a year, ~6.5% decomposition of the drug was observed. 
The plasma area under the curve (AUC) of the drug in the sub-micron emulsion 
was significantly greater than that of free drug. Overall, the drug in the emulsion 
had a lower acute toxicity and greater potential antitumor effects than the free drug, 
suggesting that the formulation is a useful tumor-targeting sub-micron emulsion 
drug delivery system. 
In a follow-up study, Junping et al.84 prepared microemulsions of vincristine 
suitable for injection using vitamin E, PEG2000-DSPE and cholesterol, adding oleic 
acid to it. The weight ratio of components used was I part drug, 5 parts oleic acid, 
5 parts vitamin E, 5 parts cholesterol and 5 parts PEG2000-DSPE. No homogenization 
was used in the preparation of the microemulsion which yielded microemulsion 
droplets of 138.1 ± 1.2 nm, when prepared using saline at pH 7.4. Note that 10 mL 
of microemulsion solution contained 1 mg of drug. The adjustment the pH of the 
aqueous phase pH and the presence of oleic acid was essential for a high drug 
loading (94.3 ± 0.3%), while the vitamin E was required for long-term storage of 
the formulation at 7 to 8°C. The formulation was stable, with respect to particle 
size, when stored at 78°C in the dark for 1 year, while the loading efficiency of 
drug decreased by approximately 3%, and 7.4% decomposition of the drug was 
observed. The plasma AUC of the vincristine in the microemulsion was significantly 
148 Lawrence & Warisnoicharoen 
greater than that of free drug. As with the previous formulation, the drug in the 
microemulsion exhibited a low acute toxicity and a high potential antitumor effect. 
4.3.2. Targeted delivery 
Shiokawa et al.M recently reported the development of a novel, tumor targeted 
microemulsion formulation suitable for delivery of the lipophilic antitumor antibiotic, 
aclacinomycin A. Tumor targeting was achieved via folate linked to the exterior 
surface of long circulating (pegylated) microemulsions. Folate was selected 
because the folate receptor is abundantly expressed in a large percentage of human 
tumors, but it is only minimally distributed in normal tissues. The basic composition 
of the microemulsion was PEG2ooo-DSPE/cholesterol/vitarnin E/drug 
(present at a 3:3:3:1 weight ratio or 7:48.3:43,3:1.5 molar ratio). In one microemulsion, 
0.24 mol% of folate linked PEG2000-DSPE was present, another contained 0.24 mol% 
of folate linked PEG5000-DSPE. In a third, the folate was linked directly to the 
DSPC and in the final one, no folate was present. The association of the folate- 
PEGsooo-linked microemulsion and folate-PEGaooo-lhiked microemulsion with the 
target cells was 200-and 4-fold higher, whereas their cytotoxicity was 90- and 3.5- 
fold higher than those of nonfolate microemulsion respectively. The folate-PEGsooolinked 
microemulsions showed 2.6-fold higher accumulation in solid tumors 24 hrs 
after i.v. injection and greater tumor growth inhibition than free drug. These findings 
suggest that a folate-linked microemulsion is a feasible means for tumortargeted 
delivery of lipophilic drug. This study shows that folate modification with 
a sufficiently long PEG chain on the exterior of a microemulsions is an effective 
way of targeting the carrier to tumor cells. 
4.4. Topical delivery 
AAA. Dermal and transdermal delivery 
The dermal and transdermal routes of administration offer several advantages compared 
with other routes of administration. However, the poor permeability of the 
stratum corneum often limits the possibilities for choosing the topical administration 
route. Therefore, novel innovative formulations such as microemulsions that 
have the potential to facilitate skin permeation are of great interest. The investigation 
of microemulsions as vehicles for cutaneous drug delivery is increasingly 
common as their potential is realized. Indeed, the cutaneous route is currently the 
most popular route of adminstration for a microemulsion. Microemulsions offer 
significant potentials as transdermal delivery vehicles, since they are robust, frequently 
stable to the addition of significant amounts of soluble enhancers, excipients 
and depending on their molecular architecture. Kreilgaard has reviewed the use of 
Recent Advances in Microemulsions as Drug Delivery Vehicles 149 
microemulsions as cutaneous drug delivery vehicles in 2002. In the present review, 
work prior to 2002 will not be dealt with in any detail. In addition, due to the large 
amount of research in the area, the review is not exhaustive. 
Proteins and peptides 
Recently, the transdermal route has received attention as a promising means to 
enhance the delivery of drug molecules, particularly peptides, across the skin, using 
harsh physical enhancement techniques such as iontophoresis and sonophoresis. 
Very little research has been performed, investigating microemulsions as vehicles 
for peptide delivery. Getie et al.66 examined the skin penetration profiles of 0.75 wt% 
desmopressin acetate released from a water-in-oil microemulsion comprising 5 wt% 
water, 20wt% Tagot 02:Span 80 3:2 and 74.25 wt% isopropyl myristate. However, 
the profile was comparable to that obtained using a standard amphiphilic cream. 
Although the amount of drug that penetrated the upper layers of the skin was 
significantly higher from the cream than from the microemulsion at all time intervals, 
within 6 hrs 6% of the applied dose reached the acceptor compartment from 
the microemulsion instead of 2% from the cream within 300 min, suggesting that 
the water-in-oil microemulsion has potential for the systemic administration of the 
drug. 
Hydrophilic drugs 
Water-in-oil microemulsions have been used to enhance the penetration of watersoluble 
drugs. For example, Alvarez-Figueroa and Blanco-Mendez9 reported the 
in vitro delivery of water-soluble methotrexate from hydrogels using iontophoresis, 
and passively from oil-in-water and water-in-oil microemulsions prepared using 
either a 3:1 v:v Labrasol: Plurol Isostearique mixture or a 3:1:1.2 v:v:v Tween 80:Span 
80:l,2-octanediol mixture as surfactant/cosurfactant, and either ethyl oleate or isopropyl 
myristate as oil. All microemulsion formulations studied were more effective 
than passive delivery from aqueous solution of the hydrophilic drug, although for 
the microemulsions, delivery was greater from the oil-in-water systems. However, 
delivery from the microemulsions was less than that using iontophoresis, probably 
because of the lower solubility of drug in microemulsions than in simple aqueous 
solution. 
Escribano et al.53 attempted to improve the transdermal permeation of sodium 
diclofenac. Four formulations were studied. One was an oil-in-water microemulsion 
based on transcutol (19wt%), plurol oleique (19.5 wt%), water (30.6 wt%), 
isostearyl isostearate (10.9 wt%) and Labrasol (19wt%). The other three formulations 
were "co-solvent" systems prepared from various of the ingredients used for 
150 Lawrence & Warisnoicharoen 
the microemulsion formulation. In this study, the microemulsion performed less 
well than the various co-solvent formulations and in a similar manner to an aqueous 
solution of the drug. This observation is perhaps not surprising as various 
enhancers were involved in the microemulsion droplets and were not available to 
improve drug penetration. Also, as it is likely that the drug was predominately in 
the continuous phase of the microemulsion, it is not surprising that the formulation 
behaved in a similar manner to an aqueous solution. 
The in vitro transdermal permeation of the antineoplastic, 5-fluorouracil, 
incorporated at I.25mg/mL in water-in-oil microemulsions prepared using 
AOT/water/isopropylmyristate has been studied by Gupta et al.69 These 
researchers found that as the water content increased from 0.9, 1.8, 2.7 and 3.6% 
w/w, microemulsions prepared with a surfactant to oil ratio of 5:95 showed 1.68, 
2.36, 3.58 and 3.77-fold increases respectively in the skin flux of 5-fluorouracil, 
compared with an aqueous solution of drug. Increasing the surfactant: oil weight 
ratio from 5:95 through 9:91 to 13:87, at fixed water:surfactant content of 15, gave 
3.58-, 5.04- and 6.3-fold enhancements of drug flux. In their study69 used attenuated 
total reflectance-Fourier transform infrared spectroscopy to determine that the 
microemulsions exerted their enhancement by interacting and perturbing the architecture 
of the statun corneum. The extent of this perturbation was dependent upon 
the concentrations of water and AOT in the microemulsion. Preliminary toxicity 
studies suggested that the microemulsions were a suitable vehicle for transdermal 
delivery. 
Amphiphilic drugs 
Jurkovic et al.85 have investigated the formulation of the amphiphilic antioxidant 
ascorbyl palmitate in a microemulsion, with a view to using the formulation as a 
protectant against free radical formation due to UV irradiation. Both oil-in-water 
and water-in-oil microemulsions were prepared using a medium chain triglyceride 
as oil, and PEG-8 caprylic/capric glycerides (Labrasol) and polyglyceryl-6-dioleate, 
(Plurol oleique) as surfactant and cosurfactant. The ascorbyl palmitate was incorporated 
into the microemulsions at various concentrations between 0.5-5.0 wt%. The 
microemulsions were gelled using either xanthan gum (water-in-oil) or Aerosil 200 
(water-in-oil). The effectiveness of the ascorbyl palmitate in the microemulsions 
depended on both the concentration and type of microemulsion. Regardless of 
the type of microemulsion, efficacy was significantly higher at the higher ascorbyl 
palmitate concentrations. Overall, the oil-in-water microemulsions were more 
effective at protecting against UV irradiation, although they delivered ascorbyl 
palmitate to the skin at a slower rate than the water-in-oil microemulsions. 
The effect of formulation composition on the in vitro release rate of the 
amphiphilic drug, diclofenac diethylamine, from a range of microemulsion vehicles 
Recent Advances in Microemulsions as Drug Delivery Vehicles 151 
containing PEG-8 caprylic/capric glycerides (surfactant), polyglyceryl-6 dioleate 
(cosurfactant), isopropyl myristate and water was determined by Djordjevic.49 The 
phase behavior of the microemulsions was determined in the absence of drug. In the 
microemulsions selected for further study, the level of water present ranged from 10 
to 60 wt% while the amount of oil varied from 8 to 46.6 wt%. The physico-chemical 
characterization studies indicated the microstructure to be either bicontinuous or 
non-spherical, and despite its amphiphilic nature, the drug was partitioned mainly 
in the water phase. The non-linearity of the drug release profile from the bicontinuous 
microemulsions was thought to be due to a complex distribution of drug 
within the microemulsion. The flux of the drug increased by >4 times, from a waterin-
oil to an oil-in-water microemulsion, the release of drug from the bicontinuous 
microemulsion, suggesting that the microstructure hampers the release of the drug. 
Hydrophobic drugs 
Dalmora and Oliveria43 and Dalmora et al.,u investigated the release of piroxicam 
encapsulated in /8-CD in cationic oil-in-water microemulsions, in an attempt to 
optimize the drug's delivery. The results demonstrated the potential of the reservoir 
in vivo system following the use of a microemulsion. The high degree of retention 
of the active substance can provide a means for modulating the anti-inflammatory 
effect, by greatly extending the release period relative to those formulations where 
the piroxicam is only dissolved or dispersed in a homogeneous aqueous medium. In 
conclusion, both microemulsions and ^-CD-containing microemulsions can offer 
many promising features for their use as topical vehicles for piroxicam delivery. 
Some of the microemulsions gelled using carbopol 940. 
Paolino et alP7 examined the potential of oil-in-water microemulsions as topical 
drug vehicles for the percutaneous delivery of ketoprofen. Microemulsions were 
prepared using triglycerides as oil, and were stabilized by a mixture of lecithin and 
n-butanol as a surfactant/ co-surfactant system. The percutaneous enhancer, oleic 
acid, was added to some of the microemulsions. Physicochemical characterization 
of the microemulsions yielded a mean droplet size of 35 nm and a negative zeta 
potential of -19.7 mV in the absence of oleic acid and — 39.5 mV in its presence. 
The ketoprofen-loaded microemulsions showed an enhanced permeation through 
excised human skin with respect to conventional formulations, although no significant 
percutaneous enhancer effect was observed in the presence of oleic acid. 
Microemulsions showed a good human skin tolerability on volunteers. 
Shukla et al.165 have investigated the potential of oil-in-water (o/w) microemulsions 
as vehicles for the dermal delivery of a eutectic mixture of lidocaine 
(lignocaine) and prilocaine, which acted as the oil phase. The microemulsion was 
stabilized by a blend of a 2:3 ratio Tween 80 and Poloxamer 331, a mixture of water 
152 Lawrence & Warisnoicharoen 
and propylene glycol were used as the hydrophilic phase. These microemulsions 
were able to solubilize up to 20 wt% of the eutectic mixture. 
In an attempt to enhance the transdermal delivery of the poorly water soluble 
drug, triptolide, and to reduce the toxicity problems associated with its usage, a 
water-in-oil microemulsion was compared with that of solid lipid nanopartides.124 
The microemulsion which was formulated using 40wt% isopropyl myristate, 
50 wt% Tween-80:l,2-propylene glycol (5:1, v/v) and water and contained 0.025 wt% 
of triptolide, gave a steady-state flux (for over 12 hours) and a permeability coefficient 
of triptolide of 6.4 ± 0.7 mg/cm2 per h and 0.0256 ± 0.002 cm/h; a value which 
was approximately double that of the solid liquid nanoparticles and 7 times higher 
than that of triptolide solution of the same concentration. In another study, Chen 
etal.33 also studied the incorporation of the drug, into a similar microemulsion using 
oleic acid as oil. Oleic acid was added because it is a known penetration enhancer, 
although there was no evidence of it acting as such in the present formulation. The 
addition, however, of 1 wt% menthol to the formulation slightly increased penetration 
from 1.58 ± 0.04 to 2.08 ± 0.06 \ig/cm2 per h (p < 0.05). Encouragingly, no 
obvious skin irritation was observed for the formulation studied, suggesting that 
microemulsions are promising vehicles for the transdermal delivery of triptolide. 
Ross et al.153 examined the transdermal penetration, across full thickness hairless 
mouse skin, of the insect repellant, N,N-diethyl-m-toluamide (DEET), contained 
in either a 1:1 v / v ethanohwater solution (containing 20 wt% DEET) or one 
of two commercially available microemulsion formulations (3M Ultrathon Insect 
Repellant (containing 31.6 wt% DEET; 3M, St. Paul, MN), and Sawyer Controlled 
Release DEET Formula (19.0%; Sawyer Products, Safety Harbor, FL). Both formulations 
were of interest because they were marketed as retarding the absorption of 
DEET due to being microemulsions. All of the DEET preparations exhibited considerable 
penetration, e.g., the ethanolic DEET formulation had a time to detection 
of approximately 30 min with steady stale at 85 min. The penetration obtained with 
the Sawyer was no different from that obtained from the ethanolic solution. The 
other microemulsion formulation (3M) demonstrated a different profile; despite 
being a higher concentration of DEET (30wt% versus 20wt%) and a comparable 
time to detection (40 min), the time to reach steady state was delayed, although 
there was still substantial absorption at steady state. 
Sintov and Shapiro168 prepared a high surfactant lidocaine microemulsion, containing 
as surfactant a mixture of glyceryl oleate and either PEG-40 stearate or 
PEG-40 hydrogenated castor oil, isopropyl myristate as oil, tetraglycol as cosurfactant, 
water, and up to 10wt% of drug, although 2.5 wt% was generally used. 
The microstructure of the microemulsion went from oil-in-water, through bicontinuous 
to water-in-oil. The penetration of the drug from the various formulations 
showed that the surfactant mixture containing PEG-40 stearate was best, while the 
Recent Advances in Microemulsions as Drug Delivery Vehicles 153 
water and surfactant/cosurfactant concentration was also important. Significantly, 
the lag time for penetration was reduced, suggesting that these microemulsions 
loaded with drug would provide rapid local analgesia. 
Priano et tzl.U7 investigated the delivery from a water-in-oil microemulsion, of 
apomorphine present as ion-pair complex with octanoate to increase its lipophilicity 
and to diminish its dissociation. As the drug was present at a high concentration, 
the dispersed phase acted as a reservoir, making it possible to maintain an almost 
constant concentration in the continuous phase and therefore achieving pseudozero-
order release kinetics. The composition of the microemulsion was complex, 
containing 18.2 wt% water, 42.1 wt% of oily phase of isopropyl miristate-decanol 
1:1.5 v/v, 3.9 wt% R-apomorphine hydrochloride, 7.3 wt% Epikuron 200, 7.1 wt% 
benzyl alcohol, 4.6 wt% octanoic acid 3.5 wt% sodium octanoate, 5.7 wt% sodium 
taurocholate, 7.6 wt% 1,2-propanediol. The microemulsion was thickened by the 
addition of 5.9 wt% Aerosil 2000. The microemulsion was able to provide in vitro, 
through hairless mouse skin, a flux of 88g/h per cm2 for 24hrs, with a kinetic 
release of pseudo-zero-order, and was chosen for in vivo study; all the components 
were biocompatible and safe. The flux gave a first approximation of the feasibility 
of the transdermal administration in man. 
The pain and discomfort caused by the injection of local anesthetics has stimulated 
research into developing topical anesthetics. However, the currently available 
formulations, such as Ametop®gel, (4 wt% amethocaine base preparation) have a 
number of disadvantages, in particular a long delay of typically 40-60 min between 
application and anesthetic effect and the requirement for a plastic occlusive dressing. 
Arevalo et alP have recently developed a decane-in-water microemulsion stabilized 
by lauromacrogol 300 and containing 4 wt% of amethocaine in an attempt 
to achieve faster drug permeation, thus reducing the time to reach optimum anesthetic 
effect. The amethocaine microemulsion proved to be a promising fast-acting 
analgesic in experimental preclinical studies. 
Mixtures of hydrophilic and hydrophobic drugs 
Although microemulsions have long been suggested as suitable formulations for 
the co-adminstration of drugs of very varying physico-chemical properties, it is only 
very recently that anyone has reported doing so. Lee et al.lw have developed a novel 
microemulsion enhancer formulation for the co-administration of hydrophilic (lidocaine 
HC1, diltiazem HC1) and lipophilic (lidocaine free base, estradiol) drugs. The 
microemulsions composed of isopropyl myristate and water, and were stabilized by 
the nonionic surfactant, Tween 80. Transdermal enhancers such as w-methyl pyrrolidone 
(NMP) and oleyl alcohol were incorporated into all systems without apparent 
disruption of the system. Unfortunately, the authors did not give the precise, 
154 Lawrence & Warisnoicharoen 
composition of the microemulsions tested; it was only mentioned that they contained 
a 1:1 v:v mixture of water and ethanol, isopropylmyristate as oil and Tween 
80 as surfactant, and were either oil-in-water or water-in-oil. Interestingly, regardless 
of the physico-chemical nature of the drug, the oil-in-water microemulsions 
provided significantly better flux for all drugs studied (p < 0.025). Enhancement 
of drug permeability from the oil-in-water systems was 17-fold for lidocaine base, 
30-fold for lidocaine HC1,58-fold for estradiol, and 520-fold for diltiazem HC1. Significantly, 
the simultaneous delivery of estradiol with diltiazem hydrochloride did 
not affect the transport of either drug (p > 0.5). 
Immunization 
Traditionally, vaccines have been administered by injection using needles, although 
the concept of topical immunization through intact skin has attracted much attention. 
Cui et al.42 recently hypothesized that a fluorocarbon-based microemulsion 
system could be one possible way to deliver plasmid DNA across the skin. 
Cui et al.42 screened a range of fluorosurfactants for their ability to form ethanolin-
perfluorooctyl bromide microemulsions. Note that the authors provided no 
evidence of a microemulsion being formed. The stability of plasmid DNA in the 
formulations was also examined. From the surfactant screen, the commercially 
available Zonyl® FSN-100, an ethoxylated nonionic fluorosurfactant, was selected 
for further study. Significant enhancements in luciferase expression and antibody 
and T-helper type-1 based immune responses, relative to those of "naked" pDNA 
in saline or ethanol, were observed after topical application of plasmid DNA in 
ethanol-in-perfluorooctyl bromide microemulsion system. From these studies, it 
can be concluded that fluorocarbon-based microemulsions are suitable for DNA 
vaccine delivery, although the mechanism(s) of the immune response induction is 
not known. It is possible that the transport of the molecules across the skin is via the 
hair-follicles, because DNA is too large and highly charged to cross intact stratum 
corneum. 
4.5. Ophthalmic 
Conventional ophthalmic dosage forms tend to be either simple solutions of watersoluble 
drugs or suspension or ointment formulations of water-insoluble drugs. 
Unfortunately, as these delivery vehicles generally result in poor levels of drug 
absorption across the cornea, most of the applied drug does not reach its intended 
site of action. However, because of the relative safety and convenience of topical 
application in ophthalmology, as well as the relatively low risk (compared 
with other routes of administration) of systemic side-effects, topical administration 
Recent Advances in Microemulsions as Drug Delivery Vehicles 1 55 
of ophthalmic agents is the preferred route of delivery. Microemulsions and submicroemulsions 
should offer a possible solution to the problem of poor delivery 
to the cornea, by sustaining the release of the drug, as well as by providing a 
higher penetration of drug into the deeper layers of the eye. In addition, they offer 
the potential of increasing the solubility of the drug in the ophthalmic delivery 
vehicle.162 
Gallarate et al.6} were probably the first to examine the potential of microemulsions 
as vehicles for ophthalmic delivery. Since then, a number of groups have 
successfully demonstrated the ability of microemulsions (sub-microemulsions) to 
prolong the ocular delivery of drug. In their study, Gallarte et al.a were able to 
further prolong the release of timolol by forming an ion pair with octanoic acid. 
Garty and Lusky63 demonstrated that the delivery of pilocarpine from an oil-inwater 
microemulsion was delayed to such an extent that the instillations of the 
microemulsion formulation twice daily were equivalent to four times daily the 
applications of conventional eye drops. A similar result was reported by Muchtar 
et alP° who determined in vitro that the corneal penetration of indomethacin formulated 
in a sub-microemulsion was more than three times that obtained using 
commercially available drops. A number of researcher have investigated the potential 
of positively charged microemulsions to retain the delivery vehicle in the eye, 
thereby sustaining delivery23,52 
To date, a range of drugs have been formulated in a microemulsion in an attempt 
to sustain release including adaprolol maleate,11'125 timolol,61 levobunolol,62 
chloramphenicol162 tepoxalin,54 piroxicam,100 delta-8-tetrahydrocannabinol,129 
pilocarpine,21'52'63'71,133 indomethacin,130 antibodies20 and dietary iso-flavonoids 
and flavonoids.83 In general, these studies showed that it was possible to delay the 
effect of drug incorporated in a microemulsion, thereby improving bioavailability. 
The proposed mechanism of the delayed action is that microemulsion droplets are 
not eliminated by the lachrymal drainage, thereby acting as drug reservoirs. The 
first studies conducted on man with microemulsions containing adaprolol maleate 
and pilocarpine, confirmed the results of the earlier studies performed mainly using 
rabbits.18,178 Vandamme178 has recently reviewed the use of microemulsions as ocular 
delivery system, and thus only studies since then will be considered in the 
present review. 
Fialho and da Silva-Cunha58 recently investigated the long term application 
of a microemulsion system in rabbits intended for the topical ocular administration 
of dexamethasone. The formulation contained 5 wt% isopropyl myristate as 
oil, 15 wt% Cremophor EL as surfactant, and a polar phase of water and 15 wt% 
propylene glycol, with dexamethasone present at a concentration of 0.1 wt%. 
Significantly, ocular irritation tests in rabbits suggested that the microemulsion did 
not provide significant alteration to eyelids, conjunctiva, cornea and iris over a Fe 
156 Lawrence & Warisnoicharoen 
3-month period. In addition, the formulation exhibited greater penetration of dexamethasone 
in the anterior segment of the eye and longer release of the drug when 
compared with a conventional preparation. The area under the curve obtained 
for the microemulsion system was more than two-fold that of the conventional 
preparation (p < 0.05). 
Gulsen and Chauhan68 have recently developed a disposable soft contact lens of 
a drug-containing microemulsion dispersed in a poly 2-hydroxyethyl methacrylate 
(HEMA) hydrogel, suitable for ophthalmic delivery, in an attempt to reduce drug 
loss and side-effects. Upon insertion into the eye, the lens will slowly release the 
drug into the pre lens (the film between the air and the lens) and the post-lens (the 
film between the cornea and the lens) tear films, thus providing a sustained delivery 
of drug. Assuming the size and drug loading of the microemulsions is low, the lenses 
should be transparent. It was found using these microemulsion-containing lenses, 
with and without a stabilizing silica shell, that drug could be released for a period 
of >8 days. By altering droplet size and loading, it is possible to tailor release. 
4.6. Vaginal 
In their 2001 review, D'Cruz and Uckun proposed that microemulsion gel formulations 
had great potential as intra vaginal/ rectal drug delivery vehicles for lipophilic 
drugs, such as microbicides, steroids, and hormones, because of their high drug 
solubilization capacity, increased absorption, and improved clinical potency, as 
long as a non toxic formulation could be prepared. In their review, D'Cruz and 
Uckun reported the formulation of two microemulsion-based gels using commonally 
available pharmaceutical excipients. Repeated intravaginal applications of formulations 
to rabbits and mice were found to be safe and did not cause local, 
systemic, or reproductive toxicity. D'Cruz and Uckun investigated the potential 
of the microemulsion-based gels as delivery vehicles of two lipophilic drugs, WHI- 
05 and WHI-07, which exhibit potent anti-HIV and contraceptive activity. As AIDS 
is spread largely through sexual intercourse, the development of a dual action 
vaginal spermicidal microbicide to curb mucosal viral transmission, as well as to 
provide fertility control would have a tremendous impact world wide. D'Cruz 
and Uckun46"48 investigated the formulation of 2 wt% of the lipophilic drugs in a 
microemulsion-based gel, composed of Phospholipon 90G and Captex 300 as the 
oil phase, with Pluronic F68 and Cremophor EL as surfactants, and seaspan carragennan 
and Xantral as gelling agents. The microemulsions were gelled to obtain the 
necessary viscosity for the gel-microemulsion formulation. Under the conditions 
of their intended use, intravaginal application of the gel-microemulsions containing 
2 wt% of drug in a rabbit model resulted in marked contraceptive activity, as 
well as exhibiting a lack of toxicity. Therefore, as a result of its dual anti-HIV and 
Recent Advances in Microemulsions as Drug Delivery Vehicles 157 
spermicidal activities, the drug-containing gels shows unique clinical potential as 
a vaginal prophylactic contraceptive for women who are at a high risk of acquiring 
HIV by heterosexual transmission. 
4.7. Nasal 
Nasal route has been demonstrated as being a possible alternative to the intravenous 
route for the systemic delivery of drugs. In addition to rapid absorption and 
avoidance of hepatic first-pass metabolism, the nasal route allows the preferential 
delivery of drug to the brain via the olfactory region, and is therefore a promising 
approach for the rapid-onset delivery of CMS medications. The solution-like feature 
of microemulsions could provide advantages over emulsions in terms of the 
sprayability, dose uniformity and formulation physical stability. 
Li et al.lu developed a diazepam-containing ethyl laurate-in-water microemulsion, 
stabilized by Tween 80 and containing propylene glycol and ethanol as cosolvents 
for the rapid-onset intranasal delivery of diazepam. A single isotropic region, 
which was considered to be a bicontinuous microemulsion, was seen at high surfactant 
concentrations but at various Tween 80: propylene glycol: ethanol ratios. 
Increasing Tween 80 concentration increased the microemulsion area, microemulsion 
viscosity, and the amount of water and oil solubilized. In contrast, increasing 
ethanol concentration produced the opposite effect. A microemulsion consisting of 
15 wt% ethyl laurate, 15 wt% water and 70 wt% Tween 80:propylene glycohethanol 
at a 1:1:1 weight ratio contained 41 mg/mL of the poorly-water soluble diazepam. 
The nasal absorption of diazepam from the formulation was fairly rapid with a maximum 
drug plasma concentration being obtained within 2 to 3 min, while bioavailability 
at 2hrs post-administration was ~50% of that obtained with intravenous 
injection. 
Zhang et al.192 attempted to prepare an oil-in-water microemulsion, containing a 
high concentration of nimodipine, suitable for brain uptake via the intranasal route 
of delivery. Three microemulsion systems stabilized by either Cremophor RH 40 or 
Labrasol, and containing a variety of oils, namely isopropyl myristate, Labrafil M 
1944CS and Maisine 35-1, were developed and characterized. The nasal absorption 
of the drug from the three microemulsions was studied in rats. The formulation composed 
of 8 wt% Labrafil M 1944CS, 30 wt% Cremophor RH 40/ethanol (3:1 weight 
ratio) and water solubilized up to 6.4 mg/mL of drug and exhibited no ciliotoxicity. 
After intranasal administration, the peak plasma concentration was obtained 
of 1 hr, while the absolute bioavailability was ~32%. Significantly, uptake of the 
drug in the olfactory bulb after nasal administration was three times that which 
was obtained from intravenous injection. In addition, the ratios of the AUC in brain 
tissues and cerebrospinal fluid to that in plasma obtained after nasal administration 
1 58 Lawrence & Warisnoicharoen 
were significantly higher than those seen after administration. In conclusion, the 
microemulsion system appears to be a promising approach for the intranasal delivery 
of nimodipine. 
Richter and Keipert51 investigated the in vitro permeability of the highly 
lipophilic material, androstenedione, across excised bovine nasal mucosa, porcine 
cornea and an artificial cellulose membrane. In order to control release, the 
two microemulsion formulations studied contained either hydroxypropyl-yScyclodextrin 
or propylene glycol. Both microemulsions were prepared from 5 wt% 
isopropyl myristate, 20 wt% Cremophor EL and water. The permeation of the drug 
through the three tissues was influenced by the microemulsion. For example, the 
apparent permeability coefficients (Papp) of the drug from the microemulsions 
across nasal mucosa did not differ from the Papp of the drug contained in solution. 
In the case of the other two membranes, release from both of the microemulsion formulations 
exhibited extended time lags, so no Papp could be calculated. It seems that 
the composition of the microemulsion had a greater impact on the Papp of cornea 
than on the Papp of the other tissues. The structure of the different membranes is 
probably responsible for the observed differences in permeation. 
4.8. Pulmonary 
Emulsions and (to a far lesser extent) microemulsions have been investigated as 
vehicles for pulmonary delivery. By far, the most widely studied systems are those 
containing fluorocarbon oil and are stabilized by a (predominately) fluorinated 
surfactant. Fluorocarbon oils are of pharmaceutical interest because of their biological 
inertness and their high (and unique) ability to dissolve gas, which means 
they can support the exchange of the respiratory gases in the lungs. In addition, 
a fluorocarbon oil, namely perfluorooctylbromide, is in Phase 11:111 clinical trials 
in the United States, for the treatment of acute respiratory distress by liquid ventilation. 
It should be noted that en-large hydrocarbon surfactants are ineffective 
solubilizers in fluorocarbon-based systems. Instead, fluorocarbon surfactants are 
required. To date, fluorocarbon-based (micro)emulsions have been investigated 
for use as oil-in-water systems for in vivo oxygen delivery (blood substitutes), 
targeted systems for diagnosis and therapy, and water-in-fluorocarbon systems 
for pulmonary drug delivery.40'102 Water-in-perfluorooctylbromide microemulsions 
have been shown to deliver homogeneous and reproducible doses of a tracer (caffeine) 
using metered-dose inhalers (pMDI) pressurized with hydrofluoroalkanes 
(HFAs).27 
Lecithin-based reverse microemulsions have also been investigated as a means 
of pulmonary drug delivery.170'171 In these studies, dimethylethyleneglycol (DMEG) 
and hexane were used as models for the propellants, dimethyl ether (DME) and 
Recent Advances in Microemulsions as Drug Delivery Vehicles 159 
propane respectively. A combination of equilibrium analysis and component diffusion 
rate determination (by pulsed-field gradient [PFG]-NMR) and iodine solubilization 
experiments were used to confirm the formation of a microemulsion. 
Water soluble solutes, including selected peptides and fluorescently labeled polya„
6-[N-(2-hydroxyethyl) D,L-aspartamide] were dissolved in the microemulsions in 
a lecithin- and water-dependent manner. Experiments with DME/lecithin demonstrated 
microemulsion characteristics similar to those in the model propellant and 
produced a droplet size and a fine particle fraction suitable for pulmonary drug 
delivery. 
Patel et al.uo have prepared water-in-hydrofluorocarbon (specifically 134a) 
microemulsions using a combination of fluorinated polyoxyethylene ether surfactants 
and a short chain hydrocarbon alcohol such as ethanol. In the absence 
of a hydrocarbon alcohol, only cosolvent systems, but not microemulsions, were 
formed. Due to the high molecular weight of the fluorocarbon surfactant, large 
concentrations of fluorocarbon surfactant are required to solubilize relatively small 
amounts of water compared with comparable hydrocarbon-based surfactants. This 
has obvious implications for the pharmaceutical application of such systems. 
To date, very little on the potential of oil-in-water microemulsions for pulmonary 
drug delivery has been investigated, yet they are attractive vehicles because 
of their ability to solubilize high amounts of drug.157 
4.8.1. Antibacterials 
Al-Adham et al.6 demonstrated that microemulsion formulations have a significant 
antimicrobial action against planktonic populations of both Pseudomonas aeruginosa 
and Staphylococcus aureus (i.e. greater than a 6 log cycle loss in viability 
over a period as short as 60s). Transmission electron microscopy studies indicated 
that this activity may in part be due to significant losses in outer membrane 
structural integrity. Nevertheless, these results have implications for the potential 
use of microemulsions as antimicrobial agents against this normally intransigent 
microorganism. 
More recently, the same group6 have determined the antibiofilm activity of 
an oil-in-water microemulsion, prepared from 15wt% Tween 80, 6wt% pentanol 
and 3wt% ethyl oleate, by incubating the microemulsion with an established 
biofilm culture of Ps. aeruginosa PA01 for a period of 4hrs. The planktonic MIC 
of sodium pyrithione and the planktonic and biofilm MICs of cetrimide were 
used as positive controls and a biofilm was exposed to a volume of normal sterile 
saline as a treatment (negative) control. The results showed that exposure to 
the microemulsion resulted in a three log-cycle reduction in biofilm viability, as 
compared to a one long-cycle reduction in viability observed with the positive 
1 60 Lawrence & Warisnoicharoen 
control treatments, suggesting that microemulsions are highly effective antibiofilm 
agents. 
5. Conclusion 
As can be seen, microemulsions are attractive d r u g delivery vehicles that offer much 
scope for improving drug delivery. Although microemulsions have been seriously 
studied as a delivery vehicle in the last >20 years, there are few microemulsion 
products currently on the market. Comparing microemulsions with vesicular drug 
delivery systems, it is pertinent to note that it took >25 years before vesicles were 
commercially exploited as drug delivery vehicles, and this was with the immense 
research effort expended in their study. Microemulsions have by contrast been much 
less widely studied. It is only a matter of time before more microemulsion-based 
formulations appear on the market. 
References 
1. Aboofazeli R, Mortazavi SA and Khoshnevis P (2003) In vitro release study of sodium 
salicylate from lecithin based phospholipid microemulsions. Iran } Pharm Res 95-101. 
2. Aboofazeli R, PatelN, Thomas M and Lawrence MJ (1995) Investigations into the formation 
and characterization of phospholipid microemulsions. 4. Pseudo-ternary phasediagrams 
of systems containing water-lecithin-alcohol and oil — the influence of oil. 
Int] Pharm 125:107-116. 
3. Agatonovic-Kustrin S and Alany RG (2001) Role of genetic algorithms and artificial 
neural networks in predicting the phase behavior of colloidal delivery systems. Pharm 
Res 18:1049-1055. 
4. Agatonovic-Kustrin S, Glass BD, Wisch MH and Alany RG (2003) Prediction of a stable 
microemulsion formulation for the oral delivery of a combination of antitubercular 
drugs using ANN methodology. Pharm Res 20:1760-1765. 
5. Agrawal GP, Juneja M, Agrawal S, Iain SK and Pancholi SS (2005) Preparation and 
characterization of reverse micelle based organogels of piroxicam. Pharmazie 59: 
191-193. 
6. Al-Adham ISI, Khalil E, Al-Hmoud ND, Kierans M and Collier PI (2000) Microemulsions 
are membrane-active, antimicrobial, self-preserving systems. / Appl Microbiol 
89:32-39. 
7. Al-Adham ISI, Al-Hmoud ND, Khalil E, Kierans M and Collier PI (2003) Microemulsions 
are highly effective anti-biofilm agents. Lett Appl Microbiol 36:97-100. 
8. Alany RG, Agatonovic-Kustrin S, Rades T and Tucker IG (1999) Use of artificial neural 
networks to predict quaternery phase systems from limited experimental data. / Pharm 
Biomed Anal 19:443-452. 
9. Alvarez-Figueroa MI and Blanco-Mendez } (2001) Transdermal delivery of methotrexate: 
Iontophoretic delivery from hydrogels and passive delivery from microemulsions. 
Int J Pharm 215:57-65. 
Recent Advances in Microemulsions as Drug Delivery Vehicles 1 61 
10. Amselem S and Friedman D (1998) Submicron emulsions in drug targeting and delivery. 
Benita S. (eds) Harwood Academic: Amsterdam. 
11. Anselem S, Beilin M and Garty N (1993) Submicron emulsion as ocular delivery system 
for adaprolol maleate, a soft b-blocker. Pharm Res 10(suppl):S025. 
12. Angelico R, Ceglie A, Colafemmina G, Lopez F, Murgia S, Olsson U and Palazzo G 
(2005) Biocompatible lecithin organogels: Structure and phase equilibria. Langmuir 21: 
140-148. 
13. Arevalo MI, Escribano E, Calpena A, Domenech J and Queralt J (2004) Rapid skin 
anesthesia using a new topical amethocaine formulation: a preclinical study. Anesth 
Anal 98:1407-1412. 
14. Atay NZ and Robinson BH (1999) Kinetic studies of metal ion complexation in glycerolin-
oil microemulsions. Langmuir 15:5056-5064. 
15. Attama AA, Nzekwe IT, Nnamani PO, Adikwu MU and Onugu CO (2003) The use of 
solid self-emulsifying systems in the delivery of diclofenac. Int J Pharm 262:23-28. 
16. Attwood D (1994) Microemulsions, in Colloidal Drug Delivery Systems. Kreuter J (ed.) 
Dekker: New York. 
17. Attwood D, Currie LRJ and Elworthy PH (1974) Studies of solubilized micellar solutions. 
1. Phase studies and particle-size analysis of solutions formed with nonionic 
surfactants. / Coll Interf Sci 46:249-254. 
18. Aviv H, Friedman D, Bar-Ilan A and Vered M (1996) Submicron emulsions as ocular 
drug delivery vehicles. US Patent 5496811. 
19. Bagwe RP, Kanicky JR, Palla BJ, Patanjali PK and Shah DO (2001) Improved drug 
delivery using microemulsions: Rationale, recent progress, and new horizons. Crit Rev 
Ther Drug 18:77-140. 
20. Becker MD, Kruse FE, Azzam L, Nobiling R, Reichling J and Volcker HE (1999) In vivo 
significance of ICAM-1 dependent leukocyte adhesion in early corneal angiogenesis. 
Invest Ophthalmol Vis Sci 40:612-618. 
21. Beilin M, Bar-Ilan A and Amselem S (1995) Ocular retention time of submicron emulsion 
(SME) and the miotic response to pilocarpine delivered in SME. Invest Ophthalmol Vis 
Sci 36:S166. 
22. Bello M, Colangelo D, Gasco MR, Maranetto F, Morel S, Podio V, Turco GL and Viano I 
(1994) Pertechnetate release from a water oil microemulsion and an aqueous-solution 
after subcutaneous injection in rabbits. ] Pharm Pharmacol 46:508-510. 
23. Benita S and Levy MY (1993) Submicron emulsions as colloidal drug carriers for intravenous 
administration — comprehensive physicochemical characterization. / Pharm 
Sci 82:1069-1079. 
24. Bhargava HN, Narurkar A and Lieb LM (1987) Using microemulsions for drug delivery. 
Pharmaceut Technol 11:46-54. 
25. Brime B, Moreno MA, Frutos G, Ballesteros MP and Frutos P (2002) Amphotericin B in 
oil-water lecithin-based microemulsions: Formulation and toxicity evaluation. / Pharm 
Sci 91:1178-1185. 
26. Brime B, Molero G, Frutos P and Frutos G (2004) Comparative therapeutic efficacy 
of a novel lyophilized amphotericin B lecithin-based oil-water microemulsion and 
1 62 Lawrence & Warisnoicharoen 
deoxycholate-amphotericin B in immunocompetent and neutropenic mice infected 
with Candida albicans. Eur J Pharm Sci 22:451-458. 
27. Butz N, Porte C, Courrier HM, Krafft MP and Vandamme TF (2002) Reverse waterin-
fluorocarbon emulsions for use in pressurized metered-dose inhalers containing 
hydrofluoroalkane propellants. hit} Pharm 238:257-269. 
28. Calvo P, Alonso MJ and Vila-Jato J (1996) Comparative in vitro evaluation of several 
colloidal systems, nanoparticles, nanocapsules, and nanoemulsions, as ocular drug 
carriers. } Pharm Sci 85:530-536. 
29. Carli F, Chiellini EE, Bellich B, Macchiavelli S and Cadelli G (2005) Ubidecarenone 
nanoemulsified composite systems, hit J Pharm 291:113-118. 
30. Castro D, Moreno MA and Lastres JL (1999) First-derivative spectrophotometric and 
LC determination of nifedipine in Brij 96 based oil/water/oil multiple microemulsions 
on stability studies. / Pharm Biomed Anal 26:563-572. 
31. Ceschel GC, Maffei P, Moretti MDL, Peana AT and Demontis S (1998) In vitro permeation 
through porcine buccal mucosa of Salvia sclarea L. essential oil from topical 
formulations. STP Pharm Sci 8:103-106. 
32. Chantrapornchai W, Clydesdale FM and McClements DJ (2001) Influence of relative 
refractive index on optical properties of emulsions. Food Res hit 34:827-835. 
33. Chen H, Chang X, Weng T, Zhao X, Gao Z, Yang Y, Xu H and Yang X (2004) A study of 
microemulsion systems for transdermal delivery of triptolide. / Control Rel 98:427-436. 
34. Chevalier Y and Zemb T (1990) The structure of micelles and microemulsions. Rep Prog 
Phy's 53:279-371. 
35. Chu L-W, Hsiue G-H and Lin I-N (2005) Ultra-fine Ba2Ti902o powders synthesized 
by inverse microemulsion processing and their microwave dielectric properties, / Am 
Ceram Soc. 
36. Constantinides PP (1995) Lipid microemulsions for improving drug dissolution and 
oral absorption: Physical and biopharmaceutical aspects. Pharm Res 12:1561-1572. 
37. Constantinides PP, Scalart JP, Marcello LJ, Marks G, Ellens H and Smith PL 
(1994) Formulation and intestinal absorption enhancement evaluation of wateri-in-oil 
microemulsions incorporating medium-chain glycerides. Pharm Res 11:1385-1390. 
38. Constantinides PP, Lancaster CM, Marcello J, Chiosone DC, Orner D, Hidalgo I, Smith 
PL, Sarkahian AB, Yiv SH and Owen AJ (1995) Enhanced intestinal absorption of 
an RGD peptide from water-in-oil nanoemulsions of different composition and size. 
J Control Rel 34:109-116. 
39. von Corswant C, Thorean P and Engstrom S (1998) Triglyceride-based microemulsion 
for intravenous administration of sparingly soluble substances. ] Pharm Sci 87:200-207. 
40. Courrier HM, Butz N and Vandamme TF (2002) Pulmonary drug delivery systems: 
Recent developments and prospects. Crit Rev Ther Drug 19:425-498. 
41. Courrier HA, Vandamme TF and Krafft MP (2004) Reverse water-in-fluorocarbon emulsions 
and microemulsions obtained with a fluorinated surfactant. Coll Surf Phy sicochem 
Eng Asp 244:141-148. 
42. Cui Z, Fountain W, Clark M, Jay M and Mumper RJ (2003) Novel ethanol-influorocarbon 
microemulsions for topical genetic immunization. Pharm Res 20:16-23. 
Recent Advances in Microemulsions as Drug Delivery Vehicles 163 
43. Dalmora MED and Oliveira AG (1999) Inclusion complex of piroxicam with (ficyclodextrin 
and incorporation in hexadecyltnmethylammonium bromide based 
microemulsion. Int J Pharm 184:157-164. 
44. Dalmora ME, Dalmora SL and Oliveira AG (2001) Inclusion complex of piroxicam with 
/S-cyclodextrin and incorporation in cationic microemulsion. In vitro drug release and 
in vivo topical anti-inflammatory effect. Int f Pharm 222:45-55. 
45. Danielsson I and Lindman B (1981) The definition of a microemulsion. Coll Surf 
3:391-392. 
46. D'Cruz OJ and Uckun MH (2001) Gel-microemulsions as vaginal spermicides and 
intravaginal drug delivery vehicles. Contraception 64:113-123. 
47. D'Cruz OJ and Uckun FM (2002) Pre-clinical safety evaluation of novel nucleoside 
analogue-based dual-function microbicides (WHI-05 and WHI-07). } Antimicrob Chem 
50:793-803. 
48. D'Cruz OJ and Uckun FM (2003) Contraceptive activity of a spermicidal aryl phosphate 
derivative of bromo-methoxyzidovudine (compound WHI-07) in rabbits. Fertil Sterio 
9:864-872. 
49. Djordjevic L, Primorac M and Stupar M (2005) In vitro release of diclofenac diethylamine 
from caprylocaproyl macrogolglycerides based microemulsions. Int} Pharm 296:73-79. 
50. Eccleston J (1994) Microemulsions, in Encyclopedia of Pharmaceutical Technology, Vol 9, 
Swarbrick J and Boylan JC (eds.) Marcel Dekker: New York. 
51. El-Aasser MS (1997) Polymeric Dispersions. Asua JM (ed.) Kluwer Academic Publications: 
The Netherlands. 
52. Elbaz E, Zeevi A, Klang S and Benita S (1993) Positively charged submicron emulsions, 
a new type of colloidal drug carrier. Int f Pharm 96:R1-R6. 
53. Escribano E, Calpena AC, Queralt J, Obach R and Domenech J (2003) Assessment 
of diclofenac permeation with different formulations: Anti-inflammatory study of a 
selected formula. Eur f Pharm Sci 19:203-210. 
54. Evitts DP, Olejnik O, Musson DG and Bilgood AM (1991) Aqueous ophtalmic 
microemulsions of tepoxalin. European patent application 0 480 690 Al. 
55. Faldt P and Bergenstahl B (1995) Fat encapsulation in spray dried food powders. / Assoc 
Offic Anal Chem 72:171-176. 
56. Fang J, Wang J, Ng S-C, Chew C-H and Gan L-M (1997) Ultrafine zirconia powders via 
microemulsion processing route. Nanostruct Mater 8:499-505. 
57. Feng L and Dexi LL (1995) Circulating emulsions (oil-in-water) as carriers for lipophilic 
drugs. Pharm Res 12:1060-1064. 
58. Fialho SL and da Silva-Cunha A (2004) New vehicle based on a microemulsion for 
topical ocular administration of dexamethasone. Clin Exp Ophthalmol 32:626-632. 
59. Florence AT and Attwood D (1998) Physicochemical Pharmacy. 3rd ed., Macmilliam 
Press, Ltd. 
60. Forgiarini A, Esquena J, Gonzalez C and Solans C (2001) Formation of nanoemulsions 
by low-energy emulsification methods at constant temperature. Eangmuir 
17:2076-2083. 
1 64 Lawrence & Warisnoicharoen 
61. Gallarate M, Gasco MR and Trotta M (1988) Influence of octanoic acid on membrane 
permeability of timolol from solutions and from microemulsions. Acta Pharm Technol 
34:102-105. 
62. Gallarate M, Gasco MR, Trotta M, Chetoni P and Saettone MF (1993) Preparation and 
evaluation in vitro of solutions and o/w microemulsions containing levobunolol as 
ion-pair. Int J Pharm 100:219-225. 
63. Garty N and Lusky M (1994) Pilocarpine in submicron emulsion formulation for treatment 
of ocular hypertension: A phase II clinical trial. Invest Ophth Vis Sci 35:2175-2179. 
64. Gasco MR, Pattarino F and Lattanzi IF (1990) Long-acting delivery systems for peptides: 
reduced plasma testosterone levels in male rats after a single injection. Int ] Pharm 
62:119-123. 
65. Gershanik T and Benita S (2000) Self-dispersing lipid formulations for improving oral 
absorption of lipophilic drugs. Eur J Pharm Biopharm 50:179-188. 
66. Getie M, Wohlrab J and Neubert RRH (2005) Dermal delivery of desmopressin acetate 
using colloidal carrier systems. / Pharm Pharmacol 57:423^127. 
67. Groves MJ and de Galindez DA (1976) The self-emulsifying action of mixed surfactants 
in oil. Acta Pharm Suet 13:361-372. 
68. Gulsen D and Chauhan A (2005) Dispersion of microemulsion drops in HEMA hydrogel: 
A potential ophthalmic drug delivery vehicle. Int f Pharm 292:95-117. 
69. Gupta RR, Jain SK and Varshney M (2005) AOT water-in-oil microemulsions as 
a penetration enhancer in transdermal drug delivery of 5-fluorouracil. Coll Surf 
Biointerf 41:25-32. 
70. Haering G and Luisi PL (1986) Hydrocarbon gels from water-in-oil microemulsions. 
JPhys Chem 90:5892-5895. 
71. Hasse A and Kiepert S (1997) Development and characterization of microemulsions for 
ocular application. Eur } Pharm Biopharm 43:170-183. 
72. He L, Wang G and Zhang Q (2003) An alternative paclitaxel microemulsion formulation: 
Hypersensitivity evaluation and pharmacokinetic profile. Int} Pharm 250:45-50. 
73. Hellweg T (2002) Phase structures of microemulsions. Curr Opin Coll Interf Sci 7:50-56. 
74. Hu Z, Tawa R, Konishi T, Shibata N and Takada K (2001) A novel emulsifier, Labrasol, 
enhances gastrointestinal absorption of gentamicin. Life Sci 69:2899-2910. 
75. Husband FA, Garrood MJ, Mackie AR, Burnett GR and Wilde PJ (2001) Adsorbed protein 
secondary and tertiary structures by circular dichroism and infrared spectroscopy 
with refractive index matched emulsions. / Agri Food Chem 49:859-866. 
76. Hwang SR, Lim S-J, Park J-S and Kim C-K (2004) Phospholipid-based microemulsion 
formulation of alMrans-retinoic acid for parenteral administration. Int ] Pharm 
276:175-183. 
77. Iek AC, Eebi NC, Tirnaksiz F and Tay A (2005) A lecithin-based microemulsion of rhinsulin 
with aprotinin for oral administration: Investigation of hypoglycemic effects in 
non-diabetic and STZ- induced diabetic rats. Int J Pharm 298:176-185. 
78. Imberg A and Engstrom S (2003) An increased throughput method for determination 
of phase diagrams/method development and validation. Coll Surf Physicochem EngAsp 
221:109-117. 
Recent Advances in Microemulsions as Drug Delivery Vehicles 165 
79. Itoh K, Matsui S, Tozuka Y, Oguchi T and Yamamoto K (2002) Improvement of physicochemical 
properties of N-4472 Part II: characterization of N-4472 microemulsion and 
the enhanced oral absorption. Int} Pharm 246:75-83. 
80. Izquierdo P, Esquena J, Tadros TF, Dederen JC, Feng J, Garcia-Celma MJ, Azemar N and 
Solans C (2004) Phase behavior and nano-emulsion formation by the phase inversion 
temperature method. Langmuir 20:6594-6598. 
81. Izquierdo P, Feng J, Esquena J, Tadros TF, Dederen JC, Garcia MJ, Azemar N and Solans 
C (2005) The influence of surfactant mixing ratio on nano-emulsion formation by the 
pit method. / Coll Interf Sci 285:388-394. 
82. Jaitely V, Sakthivel T, Magee G and Florence AT (2004) Formulation of oil in 
oil emulsions: Potential drug reservoirs for slow release. / Drug Del Sci Tech 
14:113-117. 
83. Joussen AM, Rohrschneider K, Reichling J, Kirchhof B and Kruse FE (2000) Treatment 
of corneal neovascularization with dietary iso-avonoids and flavonoids. Exp Eye Res 
71:483-487. 
84. Junping WJ, Takayama K, Nagai T and Maitani Y (2003) Pharmacokinetics and antitumor 
effects of vincristine carried by microemulsions composed of PEG-lipid, oleic 
acid, vitamin E and cholesterol. Int} Pharm 251:13-21. 
85. Jurkovic P, Sentjurc M, Gasperlin M, Kristl J and Pecar S (2003) Skin protection against 
ultraviolet induced free radicals with ascorbyl palmitate in microemulsions. Eur J Pharm 
Biopharm 56:59-66. 
86. Kanga BK, Chonb SK, Kimb SH, Jeongc SY, Kimc MS, Choc SH, Leec HB and Khanga 
G (2004) Controlled release of paclitaxel from microemulsion containing PLGA and 
evaluation of anti-tumor activity in vitro and in vivo. Int} Pharm 286:147-156. 
87. Kang LS, Jun HW and McCall JW (2000) Physicochemical studies of lidocaine-menthol 
binary systems for enhanced membrane transport. Int J Pharm 206:35-42. 
88. Kantaria S, Rees GD and Lawrence MJ (1999) Gelatin-stabilised microemulsion-based 
organogels: Rheology and application in iontophoretic transdermal drug delivery. 
/ Control Rel 60:355-365. 
89. Kantaria S, Rees GD and Lawrence MJ (2003) Formulation of electrically conducting 
microemulsion-based organogels. Int ] Pharm 250:65-83. 
90. Kawakami K, Yoshikawa T, Moroto Y, Kanaoka E, Takahashi K, Nishihara Y and 
Masuda K (2002a) Microemulsion formulation for enhanced absorption of poorly soluble 
drugs I. Prescription design. / Control Rel 81:65-74. 
91. Kawakami K, Yoshikawa T, Hayashi T, Nishihara Y and Masuda K (2002b) Microemulsion 
formulation for enhanced absorption of poorly soluble drugs II. In vivo study. 
/ Control Rel 81:75-82. 
92. Ke W-T, Lin S-Y, Ho HO and Sheu MT (2005) Physical characterizations of microemulsion 
systems using tocopheryl polyethylene glycol 1000 succinate (TPGS) as a surfactant 
for the oral delivery of protein drugs. / Control Rel 102:489-507. 
93. Khoo S-M, Porter CJH and Charman WN (2000) The formulation of halofantrine as 
either non-solubilising PEG 6000 or solubilising lipid based solid dispersions: Physical, 
stability and absolute bioavailability assessment. Int} Pharm 205:65-78. 
166 Lawrence & Warisnoicharoen 
94. Khoo S-M, Shackleford DM, Porter CJH, Edwards GA and Charman WN (2003) Intestinal 
lymphatic transport of halofantrine occurs after oral administration of a unit-dose 
lipid-based formulation to fasted dogs. Pharm Res 20:1460-1465. 
95. Khoshenvis P, Mortazavi SA, Lawrence MJ and Aboofazeli R (1997) In vitro release of 
sodium salicylate from water-in-oil microemulsions. / Pharm Pharmacol 49(S4):47. 
96. Kim C-K, Ryuu S-A, Park K-M, Lim S-J and Hwang S-J (1997) Preparation and physicochemical 
characterization of phase inverted water/oil microemulsion containing 
cyclosporin A. Int J Pharm 147:131-134. 
97. Kim C-K, Shin H-J, Yang S-G, Kim J-H and Oh Y-K (2001a) Once-a-day oral dosing 
regimen of cyclosporin A: Combined therapy of cyclosporin A premicroemulsion 
concentrates and enteric coated solid-state premicroemulsion concentrates. Pharm Res 
18:454-459. 
98. Kim C-K, Cho Y-J and Gao Z-G (2001b) Preparation and evaluation of biphenyl dimethyl 
dicarboxylate microemulsions for oral delivery. / Control Rel 70:149-155. 
99. Kim SK, Lee EH, Vaishali B, Lee S, Lee YK, Kim CY, Moon HT and Byun Y (2005) Tricaprylin 
microemulsion for oral delivery of low molecular weight heparin conjugates. 
/ Control Rel 105:32-42. 
100. Klang SH, Baszkin A and Benita S (1996) The stability of piroxicam incorporated 
in a positively charged submicron emulsion for ocular administration. Int } Pharm 
132:33-44. 
101. Kraeling MEK and Ritschel WA (1992) Development of a colonic release capsule dosage 
form and the absorption of insulin. Meth Find Exp Clin Pharmacol 14:199-209. 
102. Krafft MP, Chittofrati A and Riess JG (2003) Emulsions and microemulsions with a 
fluorocarbon phase. Curr Opin Coll Interf Sci 8:251-258. 
103. Kreilgaard M (2002) Influence of microemulsions on cutaneous drug delivery. Adv Drug 
Del Rev 54:S77-S98. 
104. Lam AC and Schechter RS (1987) The theory of diffusion in microemulsion. / Coll Interf 
Sci 120:56-63. 
105. Lawrence MJ (1994) Surfactant systems: Microemulsions and vesicles as vehicles for 
drug delivery. Eur } Drug Metab Pharmacokinet 3:257-269. 
106. Lawrence MJ (1996) Microemulsions as drug delivery vehicles. Curr Opin Coll Inter Sci 
1:826-832. 
107. Lawrence MJ and Rees GD (2000) Microemulsion-based media as novel drug delivery 
systems. Adv Drug Del Rev 45:89-121. 
108. Lee M-J, Lee M-H and Shim C-K (1995) Inverse targeting of drugs to reticuloendothelial 
system-rich organs by e:pid microemulsion emulsified with poloxamer 338. Int J Pharm 
113:175-187. 
109. Lee J-M, Park K-M, Lim S-J, Lee M-K and Kim C-K (2002) Microemulsion formulation 
of clonixic acid: Solubility enhancement and pain reduction. / Pharm Pharmacol 
54:43-49. 
110. Lee PJ, Langer R and Shastri VP (2003) Novel Microemulsion Enhancer Formulation for 
Simultaneous Transdermal Delivery of Hydrophilic and Hydrophobic Drugs. Pharm 
Res 20:264-269. 
Recent Advances in Microemulsions as Drug Delivery Vehicles 167 
111. Lettow JS, Lancaster TM, Glinka CJ and Ying JY (2005) Small-angle neutron scattering 
and theoretical investigation of poly(ethylene oxide)-poly(propylene oxide)- 
poly(ethylene oxide) stabilized oil-in-water microemulsions. Langmuir 21:5738-5746. 
112. Levy MY and Benita S (1989) Design and characterization of a submicronized o/w 
emulsion of diazepam for parenteral use. Int ] Pharm 54:103-112. 
113. Levy MY and Benita S (1991) Short-term and long-term stability assessment of a new 
injectable diazepam submicron emulsion. } Parent Sci Tech 45:101-107. 
114. Li L, Nandi I and Kim KH (2002) Development of an ethyl laurate-based microemulsion 
for rapid-onset intranasal delivery of diazepam. Int J Pharm 237:77-85. 
115. Lladser M, Medrano C and Arancibia A (1968) The use of supports in the lyophilization 
of oil-in-water emulsions. / Pharm Pharmacol 20:450-455. 
116. Lopez F, Venditti F, Ambrosone L, Colafemmina G, Ceglie A and Palazzo G (2004) 
Gelatin microemulsion-based gels with the cationic surfactant cetyltrimethylammonium 
bromide: A self-diffusion and conductivity study. Langmuir 20:9449-9452. 
117. Lu YY, Xia Q, Xia Y Ma QH and Gu N (2005) Studies on the phase behaviors of drugloading 
microemulsions. Acta Physicochimica Sinica 21:98-101. 
118. Lundberg BB (1997) A submicron lipid emulsion coated with amphipathic polyethylene 
glycol for parenteral administration of paclitaxel (Taxol). / Pharm Pharmacol 
49:16-21. 
119. Lyons KC, Charman WN, Miller R and Porter CJH (2000) Factors limiting the 
oral bioavailability of N-acetylglucosaminyl-N-acetylmuramyl dipeptide (GMDP) and 
enhancement of absorption in rats by delivery in a water-in-oil microemulsion. Int } 
Pharm 199:17-28. 
120. Malcolmson C and Lawrence MJ (1993) A comparison of the incorporation of 
model steroids into nonionic micellar and microemulsion systems. / Pharm Pharmacol 
45:141-143. 
121. Malcolmson C and Lawrence MJ (1995) Three-component non-ionic oil-in-water 
microemulsions using polyoxyethylene ether surfactants. Coll Surf B Biointerf 
4:97-109. 
122. Malcolmson C, Satra C, Kantaria S, Sidhu A and Lawrence MJ (1998) Effect of oil on the 
level of solubilization of testosterone propionate into nonionic oil-in-water microemulsions. 
/ Pharm Sci 87:109-116. 
123. Malcolmson C, Barlow DJ and Lawrence MJ (2002) Light-scattering studies of testosterone 
enanthate containing soybean oil/Ci8:iEio /water oil-in-water microemulsions. 
} Pharm Sci 91:2317-2331. 
124. Mei Z, Chen H, Weng T, Yang Y and Yang X (2003) Solid lipid nanoparticle and 
microemulsions for topical delivery of triptolide. Eur } Pharm Biopharm 56:189-196. 
125. Melamed S, Kurtz S, Greenbaum A, Haves JF, Neumann R and Garty N (1994) Adaprolol 
maleate in submicron emulsion, a novel soft /S-blocking agent, is safe and effective in 
human studies. Invest Ophthalmol Vis Sci 35:1387. 
126. Moreno MA, Ballesteros MP and Frutos P (2003) Lecithin-based oil-in-water 
microemulsions for parenteral use: Pseudoternary phase diagrams, characterization 
and toxicity studies. / Pharm Sci 92:1428-1437. 
168 Lawrence & Warisnoicharoen 
127. Mistry VV, Hassan HN and Robison DJ (1992) Effect of lactose and protein on the 
microstructure of dried milk. Food Struct 11:73-82. 
128. Myers SL and Shively ML (1992) Preparation and characterization of emulsifiable 
glasses: Oil-in-water and water-in-oil-in-water emulsions, / Coll Interf Sci 149:271-278. 
129. Muchtar S, Almong S, Torraca MT, Saettone MF and Benita S (1992) A submicron emulsion 
as ocular vehicle for delta-8- tetrahydrocannabinol: Effect on intraocular pressure 
in rabbits. Ophtalmic Res 24:142-149. 
130. Muchtar S, Abdulrazik M, Benita S (1997) Ex vivo permeation study of indomethacin 
from a submicron emulsion through albino rabbit cornea. / Control Rel 44:55-64. 
131. Nakajima H, Tomomasa S and Okabe M (1993) Preparation of Nanoemulsions, in 
Proceedings of First Emulsion Conference, Paris, EDS: Paris, Vol. 1. 
132. Nazzal S, Guven N, Reddy IK and Khan MA (2002) Preparation and characterization 
of coenzyme QlO-Eudragit solid dispersion. Drug Dev Ind Pharm 28:49-57. 
133. Naveh N, Muchtar S and Benita S (1994) Pilocarpine incorporated into a submicron 
emulsion vehicle causes an unexpectedly prolonged ocular hypotensive effect in rabbits. 
/ Ocul Pharmacol 10:509-520. 
134. Newton M, Petersson J, Podczeck F, Clarke A and Booth S (2001) The influence of 
formulation variables on the properties of pellets containing a self-emulsifying mixture. 
J Pharm Sci 90:987-995. 
135. Nyqvist-Mayer AA, Brodin AF, Frank SG (1985) Phase distribution studies on an oilwater 
emulsion based on a eutectic mixture of lidocaine and prilocaine as the dispersed 
phase. / Pharm Sci 74:1192-1195. 
136. Osborne DW, Ward AJ and O'Neill KJ (1991) Microemulsions as topical drug delivery 
vehicles: In vitro transdermal studies of a model hydroplilic drug. / Pharm Pharmacol 
43:451-454. 
137. Paolino D, Ventura CA, Nistico S, Puglisi G and Fresta M (2002) Lecithin microemulsions 
for the topical administration of ketoprofen: Percutaneous adsorption through 
human skin and in vivo human skin tolerability. Int ] Pharm 244:21-31. 
138. Park K-M and Kim C-K (1999) Preparation and evaluation of flurbiprofen-loaded 
microemulsion for parenteral delivery. Int J Pharm 181:173-179. 
139. Park K-M, Lee M-K, Hwang K-J and Kim C-K (1999) Phospholipid-based microemulsions 
of flurbiprofen by the spontaneous emulsification process. Int J Pharm 
183:145-154. 
140. Patel P, Marlow M and Lawrence MJ (2003) Formation of fluorinated nonionic surfactant 
microemulsions in hydrofluorocarbon 134a (HFC 134a). J Coll Interf Sci 258:345-353. 
141. Patil P, Joshi P and Paradkar A (2004) Effect of formulation variables on preparation 
and evaluation of gelled self-emulsifying drug delivery system (SEDDS) of ketoprofen. 
AAPS Pharm Sci Tech 5: Article 42. 
142. Pattarino F, Marengo E, Gasco MR and Carpignano R (1993) Experimental-design and 
partial least-squares in the study of complex-mixtures — microemulsions as drug carriers. 
Int} Pharm 91:157-165. 
143. Pedersen GP, Faldt P, Bergenstahl B and Kristensen HG (1998) Solid state characterisation 
of a dry emulsion: A potential drug delivery system. Int J Pharm 171:257-270. 
Recent Advances in Microemulsions as Drug Delivery Vehicles 169 
144. Podlogar F, Gasperlin M, Tomsk M, Jamnik A and Bester Rogac M (2004) Structural 
characterisation of water-Tween 40/Imwitor 308-isopropyl myristate microemulsions 
using different experimental methods. Int J Pharm 276:115-128. 
145. Porter CJH, Kaukonen AM, Boyd BJ, Edwards GA and Charman WN (2004) Susceptibility 
to lipase-mediated digestion reduces the oral bioavailability of danazol after 
administration as a medium-chain lipid-based microemulsion formulation. Pharm Res 
21:1405-1412. 
146. Pouton CW (1997) Formulation of self-emulsifying drag delivery systems. Adv Drug 
Del Rev 25:47-58. 
147. Priano L, Albani G, Brioschi A, Calderoni S, Lopiano L, Rizzone M, Cavalli R, Gasco MR, 
Scaglione F, Fraschini F, Bergamasco B and Mauro A (2004) Transdermal apomorphine 
permeation from microemulsions: A new treatment in parkinson's disease. Movement 
Disorders 19:937-942. 
148. Quellet C and Eicke H-F (1986) Mutual gelation of gelatin and water-in-oil microemulsions. 
Chimia 40:233-238. 
149. Richardson CJ, Mbanefo A, Aboofazeli R, Lawrence MJ and Barlow DJ (1997) Prediction 
of phase behavior in microemulsion systems using artificial neural networks. / Coll Inter'/ 
Sci 187:296-303. 
150. Richter A and Steiger-Trippi K (1961) Untersuchungen uber die zerstaubungstrocknung 
von emulgierten arzneizubereitungen. Pharm Acta Helv 36:322-337. 
151. Richter T and Keipert S (2004) In vitro permeation studies comparing bovine nasal 
mucosa, porcine cornea and artificial membrane: Androstenedione in microemulsions 
and their components. Eur J Pharm Biopharm 58:137-143. 
152. Ritschel WA (1991) Microemulsions for improved peptide absorption from the 
gastrointestinal-tract. Meth Findings Exp Clin Pharmacol 13:205-220. 
153. Ross EA, Savage KA, Utley LJ and Tebbett IR (2004) Insect repellant interactions: Sunscreens 
enhance deet(w, n-diethyl-m-toluamide) absorption. Drug Metabolism Dispos 
32:783-785. 
154. Rowe RC, Sheskey PJ and Weller PJ (2003) Pharmaceutical Excipients Handbook, Pharmaceutical 
Press and American Pharmaceutical Association, Dundee. 
155. Santiago RM, Fialho SL and Silva-Cunha A (2003) Design and characterization of an 
oral delivery system for insulin administration. STP Pharma Sci 13:377-380. 
156. Santos Magalhaes NS, Cave G, Seiller M and Benita S (1991) The stability and in vitro, 
release kinetics of a clofibride emulsion. Int J Pharm 76:225-237. 
157. Satra C, Thomas M and Lawrence MJ (1995) Formulating oil-in-water microemulsions 
for pulmonary drug delivery, in Drug Delivery to the Lungs IV, The Aerosol Society 
(Bristol), London. 
158. Scherlund M, Malmsten M, Holmqvist P and Brodin A (2000) Thermosetting 
microemulsions and mixed micellar solutions as drug delivery systems for periodontal 
anesthesia. Int} Pharm 194:103-116. 
159. Schwab M and Stuhn B (2000) Relaxation phenomena and development of structure in 
a physically crosslinked nonionic microemulsion studied by photon correlation spectroscopy 
and small angle x-ray scattering. / Chem Phys 112:6461-6471. 
170 Lawrence & Warisnoicharoen 
160. Schulman JH, Stoechenius W and Prince LM (1959) Mechanism of formation and structure 
of microemulsions by electron microscopy. / Phys Chem 63:1677-1680. 
161. Schurtenberger P and Cavaco C (1994) The static and dynamic structure factor of 
polymer-like lecithin reverse micelles. / Physique II 4:305-317. 
162. Siebenbrodt I and Keipert S (1993) Poloxamer-systems as potential ophthalmic 
microemulsions. Eur J Pharm Biopharm 39:25-30. 
163. Sha X, Yan G, Wu Y, Li J and Fang X (2005) Effect of self-microemulsifying drug delivery 
systems containing Labrasol on tight junctions in Caco-2 cells. Eur J Pharm Sci 
24:477-486. 
164. Shiokawa T, Hattori Y, Kawano K, Ohguchi Y, Kawakami H, Toma K and Maitani Y 
(2005) Effect of polyethylene glycol linker chain length of folate-linked microemulsions 
loading aclacinomycin A on targeting ability and antitumor effect in vitro and in vivo. 
Clin Cancer Res 11:2018-2025. 
165. Shukla A, Krause A and Neubert RRH (2003) Microemulsions as colloidal vehicle systems 
for dermal drug delivery. Part IV: Investigation of microemulsion systems based 
on a eutectic mixture of lidocaine and prilocaine as the colloidal phase by dynamic 
light scattering. / Pharm Pharmacol 55:741-748. 
166. Siebenbrodt I and Keipert S (1993) Poloxamer-systems as potential ophtalmic 
microemulsions. Eur J Pharm Biopharm 39:25-30. 
167. Sinn C (2004) When jelly gets the blues — audible sound generation with gels and its 
origin. / Non-Cryst Sol 347:11-17. 
168. Sintov AC and Shapiro L (2004) New microemulsion vehicle facilitates percutaneous 
penetration in vitro and cutaneous drug bioavailability in vivo. J Control Rel 95:173-183. 
169. Solans C, Esquena J, Forgiarini A, Uson N, Morales D, Izquierdo P, Azemar N and 
Garcia MJ (2002) Adsorption and Aggregation of Surfactants in Solution. Mittal KL and 
Shah DO, (eds.) Marcel Dekker: New York. 
170. Sommerville ML, Cain JB, Johnson CS and Hickey AJ (2000) Lecithin inverse 
microemulsions for the pulmonary delivery of polar compounds utilizing 
dimethylether and propane as propellants. Pharmaceut Dev Technol 5:219-230. 
171. Sommerville ML, Johnson CS, Cain JB, Rypacek F and Hickey AJ (2002) Lecithin 
microemulsions in dimethyl ether and propane for the generation of pharmaceutical 
aerosols containing polar solutes. Pharmaceut Dev Technol 7:273-288. 
172. Strickley RG (2004) Solubilizing excipients in oral and injectable formulations. Pharm 
Res 21:201-230. 
173. Tadros TF, Izquierdo P, Esquena J and Solans C (2004) Formation and stability of nanoemulsions. 
Adv Coll Interf Sci 108-109:303-318. 
174. Taha MO, Al-Ghazawi M, Abu-Amara H and Khalil E (2002) Development of quantitative 
structure-property relationship models for pseudoternary microemulsions formulated 
with nonionic surfactants and cosurfactants: Application of data mining and 
molecular modeling. Eur } Pharm Sci 15:461^178. 
175. Taha MO, Abu-Amara H, Al-Ghazawi M and Khalil E (2005) QSPR modeling of pseudoternary 
microemulsions formulated employing lecithin surfactants: Application of 
data mining, molecular and statistical modeling. Int} Pharm 295:135-155. 
Recent Advances in Microemulsions as Drug Delivery Vehicles 171 
176. Tenjarla S (1999) Microemulsions: An overview and pharmaceutical applications. Crit 
Rev Ther Drug Can Sys 16:461-521. 
177. Ugelstadt J, El-Aassar MS and Vanderhoff JW (1973) Emulsion polymerization — initiation 
of polymerization in monomer droplets. / Polym Sci 11:503-513. 
178. Vandamme TF (2002) Microemulsions as ocular drug delivery systems: Recent developments 
and future challenges. Prog Retin Eye Res 21:15-34. 
179. Valduga CJ, Fernandes DC, Lo Prete AC, Azevedo, CHM, Rodrigues DG and Maranhao 
RC (2003) Use of a cholesterol-rich microemulsion that binds to low-density lipoprotein 
receptors as vehicle for etoposide. / Pharm Pharmacol 55:1615-1622. 
180. Valenta C and Schultz K (2004) Influence of carrageenan on the rheology and skin 
permeation of microemulsion formulations. / Control Rel 95:257-265. 
181. Varshney M, Morey TE, Shah DO, Flint JA, Moudgil BM, Seubert CN and Dennis DM 
(2004) Pluronic microemulsions as nanoreservoirs for extraction of bupivacaine from 
normal saline. J Am Chem Soc 126:5108-5112. 
182. Vyas SP, Jain CP, Kaushik A and Dixit VK (1992) Preparation and characterisation of 
griseofulvin dry emulsion. Pharmazie 47:463-464. 
183. Wang J, Maitani Y and Takayma K (2002) Antitumor effects and pharmacokinetics of 
aclacinomycin A carried by injectable emulsions composed of vitamin E, cholesterol 
and PEG-lipid. / Pharm Sci 91:1128-1134. 
184. Wang J, Chong PP, Ng SC and Gan LM (1997) Microemulsion processing of manganese 
zinc ferrites. Mat Lett 30:217-221. 
185. Warisnoicharoen W, Lansley AB and Lawrence MJ (2000) Nonionic oil-in-water 
microemulsions: The effect of oil type on phase behaviour. Int} Pharm 98:7-27. 
186. Watnasirichaikul S, Rades T, Tucker IG and Da vies NM (2002) In vitro release and oral 
bioactivity of insulin in diabetic rats using nanocapsules,dispersed in biocompatible 
microemulsion. / Pharm Pharmacol 54:473^180. 
187. Xu QY, Nakajima M, Nabetani H, Ichikawa S and Liu XQ (2002) Factors affecting the 
properties of ethanol-in-oil emulsions. Food Sci Tech Res 8:36-41. 
188. Yamamoto A, Taniguchi T, Rikyuu K, Tsuji T, Fujita T, Murakami M and Muranishi S 
(1994) Effects of various protease inhibitors on the intestinal absorption and degradation 
of insulin in rats. Pharm Res 11:1496-1500. 
189. Yang S, Gursoy RN, Lambert G and Benita S (2004) Enhanced oral absorption of paclitaxel 
in a novel self-microemulsifying drug delivery system with or without concomitant 
use of p-glycoprotein inhibitors. Pharm Res 21:261-270. 
190. Zhang Z-Q and Lu B (2001) Advances in microemulsions as a vehicle of drug delivery 
system. Chin J Pharm 32:139-142. 
191. Zhang L, Sun X, Xiang D and Zhang ZR (2004a) Formulation and physicochemical characterization 
of norcantharidin microemulsion containing lecithin-based surfactants. 
/ Drug Del Sci Tech 14:461^69. 
192. Zhang Q, Jiang X, Jiang W, Lu, W, Su L and Shi Z (2004b) Preparation of nimodipineloaded 
microemulsion for intranasal delivery and evaluation on the targeting efficiency 
to the brain. Int J Pharm 275:85-96. 
This page is intentionally left blank
8 
Lipoproteins as Pharmaceutical Carriers 
Suwen Liu, Shining Wang and D. Robert Lu 
1. Introduction 
Large protein structures (in nanometer range) may be utilized as pharmaceutical 
carriers of drugs and DNA for targeted and other specialized delivery in biological 
systems. Lipoproteins are such structures which function as natural biological carriers 
and transport various types of lipids in blood circulation. There are many 
studies suggesting that lipoproteins can serve as efficient carriers for anticancer 
drugs, gene or other type of compounds.1-4 Previous results showed that hydrophobic 
cytotoxic drugs could be incorporated into lipoproteins, without changing the 
integrity of native lipoprotein structure. Lipoproteins as drug carriers offer several 
advantages.5-6 Firstly, they are endogenous components and do not trigger 
immunological response. They have a relatively long half-life in the circulation. Secondly, 
they have small particle size in the nanometer range, allowing the diffusion 
from vascular to extravascular compartments. Thirdly, lipoproteins can potentially 
serve as the carriers for targeted drug delivery through specific cellular receptors. 
For example, low density lipoprotein (LDL)-drug complexes may target cancer 
cells which, in many cases, have higher LDL-receptor expression than normal cells. 
Fourthly, the lipid core of lipoprotein provides a suitable compartment for carrying 
hydrophobic drugs. 
As a result of these advantages, lipoproteins have received wide attentions in 
recent years in the development of drug-targeting strategies to use them as specialized 
delivery vehicles. This review intends to provide an overview of the development 
and the specialized utilization of lipoproteins for drug delivery purpose. After 
173 
1 74 Liu, Wang & Lu 
briefly introducing the structure and the basic biological functions of lipoproteins, 
we will focus on four classes of lipoproteins, namely, chylomicron, very low-density 
lipoprotein (VLDL), low-density lipoprotein (LDL), and high-density lipoprotein 
(HDL), as the carriers for various drug compounds. Cholesterol-rich emulsions 
(LDE) and artificial lipoproteins as drug carriers will also be discussed. 
2. The Structure of Lipoproteins 
Lipoproteins, as implied by their names, are biological protein-lipid complexes. 
Lipoproteins serve the functions of carrying hydrophobic substances in blood circulation 
and transporting them to various biological sites through the protein-receptor 
interactions.6,7 The size of lipoproteins is in the nanometer range and they have a 
spherical shape with complex physicochemical properties. Figure 1 illustrates the 
general structure of lipoprotein. The hydrophobic core contains water-insoluble 
substances and is surrounded by a polar shell. The polar shell consists of phospholipids, 
unesterified cholesterol and different types of apolipoproteins, which 
bind to various cellular receptors for specific biological functions. Therefore, based 
on their physicochemical properties, lipoproteins are nanoemulsions with targeting 
functions provided by the apolipoproteins. Owing to the unique structure of 
lipoproteins, they can serve a two-mode function of solubilizing hydrophobic substances, 
including triglycerides and cholesteryl esters, within the nanoemulsion 
core and allow themselves to float in blood circulation. 
Lipoproteins can be classified into five major classes, based on their densities 
from gradient ultracentrifugation experiments. The lipoprotein classification 
includes chylomicron, very low-density lipoprotein (VLDL), intermediate-density 
lipoprotein (IDL), low-density lipoprotein (LDL), and high-density lipoprotein 
(HDL). These classes of lipoproteins have different sizes, different protein to lipid 
ratios and different types of apolipoproteins. In general, chylomicrons act on transporting 
dietary triacylglycerols and cholesterol to the adipose tissue and liver, following 
the absorption of dietary hydrophobic substances from the intestines. Very 
Fig. 1. General structure of lipoproteins. 
Lipoproteins as Pharmaceutical Carriers 1 75 
Table 1 Physicochemical properties of lipoproteins. 
Lipoprotein Transport Route Size(nm) Protein (%) Total lipids (%) 
Chylomicron Intestines to Liver 75-1200 1.5-2.5 97-99 
VLDL Liver to tissues 30-80 5-10 90-95 
IDL Liver to tissues 25-35 15-20 80-85 
LDL Liver to tissues 18-25 20-25 75-80 
HDL Tissues to liver 5-12 40-55 45-60 
low density lipoprotein, intermediate density lipoprotein and low density lipoprotein 
work at different stages to transport triacylglycerols and cholesterol from the 
liver to various tissues. High density lipoprotein brings endogenous cholesterol 
from the tissues back to the liver. The general physicochemical properties of lipoproteins 
can be seen in Table 1. 
3. Chylomicron as Pharmaceutical Carrier 
Chylomicrons are assembled in the intestine from the absorbed dietary lipids 
and transported by lymphatic system. Although most of the drugs administered 
orally are absorbed directly into the portal blood to reach the systemic circulation, 
an alternative absorption route through the intestinal lymphatics may be available 
for hydrophobic drugs. It is estimated that a high hydrophobicity (log o / w 
partition co-efficient > 5) of drug molecules is required for intestinal lymphatic 
transport.8 Chylomicrons can thus potentially serve as an important natural carrier 
for hydrophobic drugs to transport through lymphatic system.9 It is known 
that targeted drug delivery through the lymphatics is important for anti-viral drug 
molecules for the protection of B- and T-lymphocytes, which maintain relatively 
higher concentrations through the lymphatics than the systemic circulation. Chylomicrons 
have a much larger size than other lipoproteins, and thus can carry more 
drug molecules from the absorption site. With the presence of food, chylomicrons 
are the predominant lipoprotein produced by the small intestine to carry dietary 
lipids efficiently because of its large size. 
Various types of bioactive molecules have been incorporated into reconstituted 
chylomicron structure for delivery purposes. In gene delivery, Hara et a/.10,11 developed 
reconstituted chylomicron which incorporated a hydrophobic DNA complex 
and used it as an in vivo gene transfer vector. They found that the DNA-incorporated 
chylomicrons induced a high gene expression in mouse liver after the reconstituted 
chylomicron was administered through portal vain injection. Furthermore, 
it was also reported that artificial, protein-free lipid emulsions could be utilized to 
model the metabolism of lymph chylomicron in rats, not only in the initial partial 
176 Liu, Wang & Lu 
hydrolysis by lipoprotein lipase, but also in the delivery of a remnant-like particle 
to the liver.12 As a targeted therapeutic approach to hepatitis B, anti-viral iododeocyuridine 
was incorporated into recombinant chylomicrons, resulting in the drug 
molecules being selectively targeted to the liver parenchymal cells.13 It has been 
suggested that chylomicron can serve as a special carrier for liver cell targeting.14 
Due to the targetability, this approach could be further developed as an effective 
therapy for hepatitis B patients. 
4. VLDL as Pharmaceutical Carrier 
VLDL particles have a size range of 30-80 nm. They are assembled in the endoplasmic 
reticulum (ER) and matured in Golgi apparatus of hepatocytes before 
secretion.15 After entering into the plasma, VLDL particles are catabolized by a 
series of biochemical actions, including apolipoprotein exchange of apoC-I, apoCII, 
apoC-III, and apoE; lipolysis by triglyceride lipase; and cell-surface receptormediated 
uptake. As lipolysis proceeds, VLDL particles become smaller and are 
eventually converted to IDL. Some of the IDL particles are rapidly taken up by hepatocytes 
via a receptor-mediated mechanism while others undergo further hydrolysis 
before being converted to LDL. The catabolism route of VLDL suggests the 
possibility of using VLDL as a drug carrier for targeted delivery. ApoE is a protein 
ligand present on the surface of VLDL and it is well known that the receptor of 
apoE is overexpressed on some types of cancer cells. Therefore, VLDL can potentially 
serve as an antineoplastic drug carrier. 
As a drug carrier, VLDL is an interesting candidate because it contains a relatively 
small amount of proteins (about 5-10 % protein) and a large amount of triglycerides 
(about 50-65% within the emulsion core) which can be used to solubilize 
hydrophobic substances sufficiently. By mimicking the compositions and structure 
of VLDL, Shawer et al. developed a VLDL-resembling phospholipid nanoemulsion 
system that carried a new anti-tumor boron compound for targeted delivery to cancer 
cells.16 The nanoemulsion demonstrated sufficient capability to solubilize the 
hydrophobic compound. The structure of the phospholipid nanoemulsion was verified 
based on the changes in the molecular surface area and the molecular volume 
of each component of the nanoemulsion when the particle size is changed (from different 
size fractions). If certain molecules are located at the core of nanoemulsion, 
their numbers per overall volume should not be changed when the particle size 
is increased. If certain molecules are located at the surface of nanoemulsion, their 
numbers per overall volume should decrease when particle size is increased. This is 
because the overall surface area decreases when particle size is increased. Similar to 
the natural lipoproteins, it was demonstrated that phospholipid was predominately 
Lipoproteins as Pharmaceutical Carriers 177 
located at the surface and the hydrophobic substances, triolein and cholesteryl 
oleate, were mainly located in the core of the phospholipid nanoemulsion. 
Recently, a similar nanoemulsion formulation was used to encapsulated 
quantum dots (QD) as a new bioimaging carrier.17 Quantum dots (QDs) are 
semiconductor nanocrystals that are emerging as unique fluorescence probes in 
biomedicine.18-21 When manufactured, most of the quantum dots have organic ligand 
coating on their surface and are extremely hydrophobic. The research goal was 
to encapsulate QDs in phospholipid nanoemulsion and to examine the physical 
stability, size distribution and their interactions with cancer cells. It was found that 
CdSe QDs can be efficiently encapsulated in the phospholipid nanoemulsion. The 
QD-encapsulated phospholipid nanoemulsion are stable and interact well with cultured 
cells to deliver the QDs inside the cells for fluorescence imaging.17 In other 
studies, it has been demonstrated that cytotoxic drugs such as 5-fluorouracil (5-FU), 
5-iododeoxyuridine (IudR), doxorubicin (Dox), and vindesine can be effectively 
incorporated into VLDL, and the resultant complexes showed effective cytotoxicity 
to human carcinoma cells.22 
5. LDL as Pharmaceutical Carrier 
LDL (18-25 nm) is not directly synthesized in human body. Instead, most of them 
are formed through the VLDL pathway. LDL is the major circulatory lipoprotein 
for the transport of cholesterol and cholesteryl esters, and it can be internalized by 
cells via LDL receptor-mediated endocytosis. The internalization process of LDL 
has been well characterized and the understanding of the mechanism can potentially 
help the designing of the drug targeting strategy through the LDL receptor 
(Fig. 2). The binding of dephosphorylated adaptor protein to the plasma membrane 
LDL Receptors 
(. ( . X l B l O l f c . H K . * ^ . , . ^ 
Cell , HMGCoA 
t ACAT T 
Cholesterol 
\ LDL Receptors 
mug •> ^-.t, ^ v f i ' * o„o -* 
LDL Binding —• Internalization —•Drug Release —^Regulation 
Fig. 2. LDL receptor pathway and targeted drug delivery. 
1 78 Liu, Wang & Lu 
initiates the formation of coated pits which are covered by the protein clathrin. The 
receptors from the surrounding regions of the plasma membrane shift towards the 
binding site for internalization. Apolipoproteins including apo B-100 and apo E 
are recognized and bound by the LDL receptor on the cell surface to form a complex 
which is internalized into the coated pits. After internalization of the LDL, 
the coated pits are pinched off and within a very short time, they shed off their 
clathrin coating. The internalized LDL particle is transferred to endocytotic vesicles 
or endosomes. Due to the acidic pH within the endosomes, LDL dissociates 
from its receptor. This is followed by the fusion of the endosomes with lysosomes 
which contain hydrolases. The protein component of LDL is broken into free amino 
acids, while the cholesteryl ester component is cleaved by lysosomal lipase. The free 
cholesterol is released and incorporated into the cell membrane. Excess cholesterol 
is re-esterified by the action of acyl-CoA:cholesterol acyltransferase (ACAT). 
Among various lipoproteins, LDL has been widely studied as a drug carrier for 
targeted and other specialized deliveries, because many types of cancer cells show 
elevated expression of LDL receptors than the corresponding normal cells.23-26 
Comparing with chylomicron, VLDL, and IDL, LDL also has a longer serum halflife 
of 2-4 days,27 making it a desirable drug carrier. Low density lipoprotein was 
found to be suitable as carriers for cytotoxic drugs to target cancer cells. LDLdrug 
complexes can be formed through various processes without changing the 
lipoprotein integrity.28-31 
5.1. LDL as anticancer drug carriers 
Doxorubicin (Dox) is widely used in treating different tumors. Its main side effects 
are cadiotoxicity and multidrug resistance, especially during prolonged treatment 
in the patients. LDL has been studied as a target carrier for Dox in nude mice, bearing 
human hepatoma HepG2 cells.32 Both in vitro and in vivo studies indicated that when 
Dox was incorporated into LDL, the multidrug resistance could be circumvented 
and the cardiotoxicity could be reduced as well.33 Kader and Pater22 used VLDL, 
LDL and HDL as carriers to deliver four cytotoxic drugs, 5-fluorouracil (5-FU), 
5-iododeoxyuridine (IUdR), doxorubicin (Dox) and vindesine. They found that 
significant drug loading was achieved in all three classes of lipoproteins, consistent 
with the sizes and hydrophobicity of the drug. Experiments were carried out to 
examine the changes in drug cytotoxicity against HeLa cervical and MCF-7 breast 
carcinoma cells, after the incorporation into lipoprotein. The results demonstrated 
that VLDL-drug complex did not affect their IC50 on both HeLa and MCF-7 cell 
lines, when compared with free drugs. However, the IC50 values of LDL- and HDLdrug 
complexes were significantly lower compared with free drugs. Their studies 
further indicated that drugs were incorporated into lipoproteins without disrupting 
Lipoproteins as Pharmaceutical Carriers 1 79 
their integrity; drugs remained in their stable forms inside lipoproteins; and human 
LDL and HDL could be particularly useful in the delivery of antineoplastic drugs. 
5.2. LDL as carriers for other types ofbioactive compounds 
Although LDL has been widely studied as a carrier to deliver anticancer compounds, 
it may also be useful to deliver other types of bioactive compounds. 
LDL may serve as a carrier for site-specific delivery of drugs to atherosclerotic 
lesions.34 When dexamethasone palmitate (DP), a steroidal anti-inflammatory drug, 
was incorporated in LDL, an inhibitory effect of this complex on foam cell formations 
was demonstrated. The study indicated that LDL could potentially carry 
DP to atherosclerotic lesions.34 Fluorophore-labeled LDL was also used for optical 
imaging in tumors diagnosis. For example, carbocyanine dyes can be used 
as near infrared (NIR) optical imaging probes with long wavelength absorption, 
high extinction coefficients and high fluorescence quantum yield. In vitro confocal 
microscopic study and ex vivo low-temperature fluorescent scanning demonstrated 
that carbocynine-labled LDL probes, Dil-LDL, could be selectively delivered to 
B16/HepG2 tumor cells and the corresponding animal tumors via the LDL receptor 
pathway.35 It was also proposed that Dil is located and oriented in the phospholipid 
monolayer when it binds to LDL. 
5.3. LDL for gene delivery 
LDL has also been investigated as gene delivery carriers. Comparing with viral 
gene-delivery vectors and some other types of non-viral gene delivery vectors, 
the LDL system shows certain advantages in transfection efficiency and safety 
considerations.5 Several LDL based gene delivery systems have been reported. 
Kim's group developed a terplex system which comprises LDL, lipidized poly(Llysine) 
and plasmid DNA. The complex had a diameter of about 100 nm. The studies 
showed high efficiency to deliver plasmid DNA to smooth muscle cells and fibroblast 
cells.36,37 In addition, a novel LDL-DNA complex was formulated by Khan 
et al.38 LDL was cationized using carbodiimide and the modified lipoprotein complex 
significantly increased the DNA binding capacity with improved stability. The 
novel delivery system also demonstrated the ability to target cells through LDL 
receptor.38 
6. HDL as Pharmaceutical Carriers 
Among various lipoproteins, HDL has the smallest size with a diameter of 5-12 nm. 
It shares common structural characteristics with other lipoproteins. However, its 
180 Liu, Wang & Lu 
polar shell contributes more than 80% of the total mass. Newly synthesized HDL 
hardly contains any cholesteryl ester molecules. Cholesteryl esters are gradually 
added to the particles by lecithin via enzymatic reaction: cholesterol acyltransferase 
(LCAT), which is a 59-kD glycoprotein associated with HDL. The interaction of 
HDL with cells appears similar to that of LDL.39 Although the function of HDL in 
the human body is not well-defined, it generally transports excess cholesterol and 
cholesteryl esters from various tissue cells back to the liver. Comparing with other 
types of lipoproteins, small size and fast internalization by tumor cells are the major 
advantages of utilizing HDL for drug delivery and targeting. 
HDL has mainly been utilized for the delivery of water insoluble anticancer 
drugs through the targeting function.40-41 When the anticancer drug, Taxol, was 
incorporated into HDL, stable complexes were formed and they were examined for 
cancer-cell targeting.41 Reconstituted HDL was explored as a drug carrier system for 
a lipophilic prodrug, IDU-OI2.42 The studies indicated that the lipophilic prodrug 
could be efficiently incorporated into reconstituted HDL particles. This approach 
may also be useful to encapsulate other lipophilic derivatives of water-soluble 
drugs. The utilization of HDL for drug targeting may lead to a more effective therapy 
for infectious diseases, such as hepatitis B, since the HDL-drug complexes were 
demonstrated to be selectively taken by parenchymal liver cells.42 Comparing with 
free drugs in cytotoxicity assays, the IC values of HDL-drug complexes were significantly 
decreased, about 2.5 to 23-fold lower.22 Interestingly, it was observed that 
HDL-drug complex specifically increased the cytotoxicity to carcinoma cells. Earlier 
studies showed that HDL could increase the sensitivity of HeLa cells to the 
cytotoxic effects of Dox.43 Similar to LDL-drug complex, the lipoprotein receptor 
pathway appears to be involved in the interactions between HDL-drug complex 
and cancer cells. 
7. Cholesterol-rich Emulsions (LDE) as Pharmaceutical 
Carriers 
LDE is a lipid based formulation, an emulsion with a lipid structure resembling LDL 
particle and it is made without protein incorporation. Essentially, it is composed of 
a cholesteryl ester core surrounded by a monolayer of phospholipids. Comparing 
with native LDL, LDE is removed from the blood circulation more rapidly.44 It 
appears possible that LDE can acquire apoE and other apolipoproteins from native 
lipoproteins in plasma. ApoE can be recognized by LDL receptors, thus allowing the 
binding of LDE to the receptors. However, it is known that LDE binds to receptors 
through apoE, but not through apoBlOO. The interaction between apoE and the 
receptor appears stronger than that of apoBlOO.45 
Lipoproteins as Pharmaceutical Carriers 181 
LDE is considered as a potential carrier for anticancer drugs to deliver 
chemotherapeutic agents to neoplastic cells. Although there is no protein in the 
LDE formulations, previous studies showed that the LDL receptor could still play 
an important role in the cellular uptake of these lipid complexes.46-56 LDE binds 
to low-density lipoprotein receptors which are upregulated in cancer cells, leading 
to a higher concentration in neoplastic tissues.24-57 LDE-carmustine complex was 
studied with a neoplastic cell line and its biodistribution was studied in mice. An 
exploratory clinical study was also conducted. The result showed that the uptake of 
LDE-carmustine complex by tumor was several fold greater than the uptake by the 
corresponding normal tissue. The association of carmustine with LDE preserves the 
tumor-cytotoxicity of carmustine with reduced side effects.58 Preliminary clinical 
study59 was also carried out using LDE-carmustine complex to treat patients with 
advanced cancers. The results demonstrated that the systemic toxicity of the drug 
was significantly reduced. 
Rodrigues et ah investigated the formulation of LDE containing antineoplastic 
compound paclitaxel.55 The experiments revealed a 75% incorporation efficiency 
and the stable complex of the drug molecules incorporated in LDE emulsion. Its 
LD50 was ten-fold greater than that of a commercial formulation of paclitaxel. It was 
suggested by the authors that the cellular uptake and the cytotoxic activity of LDEpaclitaxel 
complex might be mediated by the LDL receptors due to the cholesterol 
moiety in the LDE formulation.55 
In addition to LDE, artificial lipoproteins have been constructed. Several 
research groups have developed various types of artificial lipoproteins.44-60-62 Most 
of them constructed the artificial lipoproteins by incorporating natural apoB protein 
into lipid microemulsion for the purpose of examining the lipoprotein metabolism. 
Artificial lipoproteins containing poly-lysine has also been investigated as the DNA 
carrier for cellular transfection, with the potential to reduce the cytotoxicity and to 
improve the transfection efficiency.63-64 
8. Concluding Remark 
Lipoproteins are natural nanostructures in biological systems. They have unique 
physicochemical properties which may be utilized as pharmaceutical carriers for 
drug compounds and other bioactive substances. Owing to the structural diversity 
of lipoproteins, including chylomicron, VLDL, LDL and HDL, various specialized 
delivery systems may be developed to fully utilize their delivery potentials. New 
nanostructures, such as LDE and artificial lipoproteins, can also be constructed to 
mimic the structure of natural lipoproteins. As these new nanostructures are built 
from scratch, they may be more efficient in encapsulating drug and other bioactive 
molecules, and more effective for specialize drug delivery. 
1 82 Liu, Wang & Lu 
References 
1. Smidt PC and van Berkel TJC (1990) LDL-mediated drug targeting. Crit Rev Ther Drug 
Carrier Syst 7:99-119. 
2. Filipowska D, Filipowski T, Morelowska B, Kazanowska W, Laudanski T, Lapinjoki S, 
Akerland M and Breeze A (1992) Treatment of cancer patients with a low density 
lipoprotein delivery vehicle containing a cytotoxic drug. Cancer Chemother Pharmacol 29: 
396-400. 
3. Firestone RA (1994) Low density lipoprotein as a vehicle for targeting antitumor compounds 
to cancer cells. Bioconjug Chem 5:105-113. 
4. van Berkel TJC (1993) Drug targeting: Application of endogenous carriers for site specific 
delivery of drug. / Control Rel 24:145-155. 
5. Pan G, 0ie S and Lu DR (2003) Biological protein nanostructures and targeted drug 
delivery. Lu DR and 0ie S (eds.) in Cellular Drug Delivery: Principles and Practice, 
pp. 217-234. 
6. Chung NS and Wasan KM (2004) Potential role of the low-density lipoprotein receptor 
family as mediators of cellular drug uptake. Adv Drug Del Rev 56:1315-1334. 
7. Sarkar R, Halpern DS, Jacobs SK and Lu DR (2002) LDL-receptor mediated drug targeting 
to malignant tumors. Muzykantov VR and Torchilin VP (eds.) in Biomedical Aspects of 
Drug Targeting. (Kluwer Academic Publisher), pp. 327-345. 
8. Charman WN and Stella VJ (1986) Estimating the maximal potential for intestinal lymphatic 
transport of lipophilic drug molecules, hit} Pharm 34:175-178. 
9. Shen H, Howies P and Tso P (2001) From interaction of lipidic vehicles with intestinal 
epithelial cell membranes to the formation and secretion of chylomicrons. Adv Drug Del 
Rev 50:S103-S125. 
10. Hara T, Liu F, Liu DX and Huang L (1997) Emulsion formulations as a vector for gene 
delivery in vitro and in vivo. Adv Drug Del Rev 24:265-271. 
11. Hara T, Tan Y and Huang L (1997) In vivo gene delivery to the liver using reconstituted 
chylomicron remnants as a novel nonviral vector. Proc Natl Acad Sci USA 94: 
14547-14552. 
12. Redgrave TG and Maranhao RC (1985) Metabolism of protein-free lipid emulsion models 
of chylomicrons in rats. Biochem Biophys Acta 835:104-112. 
13. Rensen PCN, De Vrueh RLA, van Berkel TJC (1996) Targeting hepatitis B therapy to the 
liver: Clinical pharmacokinetic considerations. Clin Pharmacokinet 31:131-155. 
14. Rensen PC, van Dijk MC, Havenaar EC, Bijsterbosch MK, Kruijt JK and van Berkel 
TJ (1995) Selective liver targeting of antivirals by recombinant chylomicrons — a new 
therapeutic approach to hepatitis B. Nat Med l(3):221-5. 
15. Olofsson SO, Bjursell G, Bostrom K, Carlsson P, Elovson J, Protter AA, Reuben MA and 
Bondjers G (1987) Apolipoprotein B: Structure, biosynthesis and role in the lipoprotein 
assembly process. Atherosclerosis 68:1-17. 
16. Shawer M, Greenspan P, 0ie S and Lu DR (2002) VLDL-resembling phospholipidsubmicron 
emulsion for cholesterol-based drug targeting. / Pharm Sci 91:1405-1413. 
Lipoproteins as Pharmaceutical Carriers 183 
17. Liu S, Lee CM, Wang S and Lu DR (2006) A new bioimaging carrier for quantum dot 
nanocrystals — phospholipid nanoemulsion mimicking natural lipoprotein core. Drug 
Del 13:159-164. 
18. Dubertret B, Skourides P, Norris DJ, Noireaux V, Brivanlou AH and Libchaber A (2002) 
In vivo imaging of quantum dots encapsulated in phospholipid micelles. Science 298: 
1759-1762. 
19. Gao X, Cui Y, Levenson RM, Chung LW and Nie S (2004) In vivo cancer targeting and 
imaging with semiconductor quantum dots. Nat Biotechnol 22:969-976. 
20. Bruchez M, Moronne M, Gin P, Weiss S and Alivisatos AP (1998) Semiconductor 
nanocrystals as fluorescent biological labs. Science 281:2013-2016. 
21. Chan WCW and Nie S (1998) Quantum dot biocojugates for ultrosensitive nonisotopic 
detection. Science 281:2016-2018. 
22. Kader A and Pater A (2002) Loading anticancer drugs into HDL as well as LDL has little 
affect on properties of complexes and enhances cytotoxicity to human carcinoma cell. 
/ Control Rel 80:29^4. 
23. Alexopoulos CG, Blatsios B and Avgerinos A (1987) Serum lipids and lipoprotein disorders 
in cancer patients. Cancer 3065-3070. 
24. Ho YK, Smith RG, Brown MS and Goldstein JL (1978) Low density lipoprotein (LDL) 
receptor activity in human acute myelogenous leukemia cells. Blood 52:1099-1114. 
25. Klock JC and Pieprzyk JK (1979) Cholesterol, phospholipids, and fatty acids of normal 
immature neutrophils: Comparison with acute myeloblastic leukemia cells and normal 
neutrophils. / Lipid Res 20:908-911. 
26. Nakagawa T, Ueyama Y, Nozaki S, Yamashita S, Menju M, Funahashi T, Takemura KK, 
Kubo M, Tokunaga K, Tanaka T, Yagi M and Matsuzawa Y (1994) Marked hypocholesterolemia 
in a case with adrenal adenoma — Enhanced Catabolism of low density 
lipoprotein (LDL) via the LDL receptors of tumor cells. / Clin Endocrinol Metabol 80: 
92-96. 
27. Kader A, Davis PJ, Kara M and Liu H (1998) Drug targeting using low density lipoprotein 
(LDL): Physicochemical factors affecting drug loading into LDL particles. / Control Rel 
55:231-243. 
28. Firestone RA (1994) Low density lipoprotein as a vehicle for targeting antitumor compounds 
to cancer cells. Bioconjug Chem 5:105-113. 
29. Filipowska D, Filipowski T, Morelowska B, Kazanowska W, T. Laudanski T, Lapinjoki S, 
Akerland M and Breeze A (1992) Treatment of cancer patients with a low density lipoprotein 
delivery vehicle containing a cytotoxic drug. Cancer Chemother Pharmacol 29:396^00. 
30. de Smidt PC and van Berkel TJC (1990) LDL-mediated drug targeting. Crit Rev Ther Drug 
Can Syst 7:99-119. 
31. van Berkel TJC (1993) Drug targeting: Application of endogenous carriers for site specific 
delivery of drug. / Control Rel 24:145-155. 
32. Chu ACY, Tsang SY, Lo EHK and Fung KP (2001) Low density lipoprotein as a targeted 
carrier for doxorubicin in nude mice bearing tumor hepatoma HepG2 cells. Life Sci 
70:591-601. 
184 Liu, Wang & Lu 
33. Lo EHK, Ooib VEL and Fung KP (2002) Circumvention of multidrug resistance and 
reduction of cardiotoxicity of doxorubicin in vivo by coupling it with low density lipoprotein. 
Life Sci 72:677-687. 
34. Tauchi Y, Takase M, Zushida I, Chono S, Sato J, Ito K and Morimoto K (1999) Preparation 
of a complex of dexamethasone palmitate-low density lipoprotein and its effect on foam 
cell formation of murine peritoneal macrophages. / Pharma Sci 88:709-714. 
35. Li H, Zhang Z, Blessington D, Nelson DS, Zhou R, Lund-Katz S, Chance B, Glickson 
JD and Zheng G (2004) Carbocyanine labeled LDL for optical imaging of tumors. Acad 
Radiol 11:669-677. 
36. Kim JS, Maruyama A, Akaike T and Kim SW (1997) Ln vitro gene expression on smooth 
muscle cells using a terplex delivery system. / Control Rel 47:51-59. 
37. Kim JS, Kim BI, Maruyama A, Akaike T and Kim SW (1998) A new non-viral DNA 
delivery vector: The terplex system. / Control Rel 53:175-182. 
38. Khan Z, O. Hawtrey A and Ariatti M (2003) New cationized LDL-DNA complexes: Their 
targeted delivery to fibroblasts in culture. Drug Del 10:213-220. 
39. Steinberg D (1996) A docking receptor for HDL cholesterol esters. Science 271: 
46CM61. 
40. Rensen PC, de Vrueh RL, Kuiper J, Bijsterbosch MK, Biessen EA and van Berkel TJ (2001) 
Recombinant lipoproteins: Lipoprotein-like lipid particles for drug targeting. Adv Drug 
Del Rev 47(2-3):251-276. 
41. Lacko AG, Nair M, Paranjape S, Johnso S and McConathy WJ (2002) High density 
lipoprotein complexes as delivery vehicles for anticancer drugs. Anticancer Res 22: 
2045-2049. 
42. Bijsterbosch MK, Schouten D and van Berkel TJ (1994) Synthesis of the dioleoyl 
derivative of iododeoxyuridine and its incorporation into reconstituted high density 
lipoprotein particles. Biochemistry 33:14073-14080. 
43. Chassany O, Urien S, Claudepierre P, Bastian G and Tillement JP (1996) Comparative 
serum protein binding of anthra- cycline derivatives. Cancer Chemother Pharmacol 38: 
571-573. 
44. Hirata RDC, Hirata MH, Mesquita CH, Cesar TB and Maranhao RC (1999) Effects of 
apolipoprotein B-100 on the metabolism of a lipid microemulsion model in rats. Biochim 
Biophys Acta 1437:53-62. 
45. Innerarity TL and Mahley RW (1978) Enhanced binding by cultured human broblasts of 
apo-E-containing lipoproteins as compared with low density lipoproteins. Biochemistry 
17:1440. 
46. Versluis AJ, Rump ET, Rensen PC, van Berkel TJ and Bijsterbosch MK (1998) Synthesis of 
a lipophilic daunorubicin derivative and its incorporation into lipidic carriers developed 
for LDL receptor-mediated tumor therapy. Pharm Res 15:531-537. 
47. Versluis AJ, Rensen PC, Rump ET, van Berkel TJ and Bijsterbosch MK (1998) Lowdensity 
lipoprotein receptor-mediated delivery of a lipophilic daunorubicin derivative 
to B16 tumours in mice using apolipoprotein E-enriched liposomes. Br J Cancer 78: 
1607-1614. 
Lipoproteins as Pharmaceutical Carriers 185 
48. Amin K, Wasan KM, Albrecht RM and Heath TD (2002) Cell association of liposomes 
with high fluid anionic phospholipids content is mediated specifically by the LDL and 
its receptor. / Pharm Sci 91:1233-1244. 
49. Amin K, Ng K, Brown CS, Bruno MS and Heath TD (2001) LDL induced association 
of anionic liposomes with cells and delivery of contents as shown by the increase in 
potency of liposome dependent drugs. Pharm Res 18:914-921. 
50. Amin K and Heath TD (2001) LDL-induced association of anionic liposomes with cells 
and delivery of contents: II. Interaction of liposomes with cells in serum-containing 
medium. / Control Rel 73:49-57. 
51. Lakkaraju A, Rahman Y and Dubinsky JM (2002) Low-density lipoprotein-related 
protein mediates the endocytosis of anionic liposomes in neurons. / Biol Chem 277: 
15085-15092. 
52. Rensen PC, Schiffelers RM, Versluis AJ, Bijsterbosch MK, Van Kuijk-Meuwissen ME 
and van Berkel TJ (1997) Human recombinant apolipoprotein E-enriched liposomes can 
mimic low-density lipoproteins as carriers for site-specific delivery of antitumor agents. 
Mol Pharmacol 52:445^55. 
53. Koller-Lucae SKM, Schott H and Schwendener RA (1997) Interactions with human 
blood in vitro and pharmacokinetic properties in mice of liposomal N4-octadecyl-l-h- 
D-arabinofuranosylcytosine, a new anticancer drag. J Pharmacol Exp Ther 282:1572-1580. 
54. Koller-Lucae SKM, Schott H and Schwendener RA (1999) Lowdensity lipoprotein 
and liposome mediated uptake and cytotoxic effect of N4-octadecyl-l-h-Darabinofuranosylcytosine 
in Daudi lymphoma cells. Br ] Cancer 80:1542-1549. 
55. Rodrigues DG, Covolan CC, Coradi ST, Barboza R and Maranhao RC (2002) Use of a 
cholesterol-rich emulsion that binds to low-density lipoprotein receptors as a vehicle for 
paclitaxel. / Pharm Pharmacol 54:765-772. 
56. Versluis AJ, Rump ET, Rensen PC, van Berkel TJ and Bijsterbosch MK (1999) Stable incorporation 
of lipophilic daunorubicin prodrug into apolipoprotein E-exposing liposomes 
induces uptake of prodrug via low-density lipoprotein receptor in vivo. J Pharmacol Exp 
Ther 289:1-7. 
57. Maranhao RC, Roland IA, Toffoletto O, Ramires JA, Gone, alves RP, Mesquita CH 
and Pileggi P (1997) Plasma kinetic behavior in hyperlipidemic subjects of a lipidic 
microemulsion that binds to LDL receptors. Lipids 32:627-633. 
58. Maranhao RC, Graziani SR, Yamaguchi N, Melo RF, Latrilha MC, Rodrigues DG, 
Couto RD, Schreier S and Buzaid AC (2002) Association of carmustine with a lipid 
emulsion: In vitro, in vivo and preliminary studies in cancer patients. Cancer Chemother 
Pharmacol 49:487-^198. 
59. Hungria VTM, Latrilha MC, Rodrigues DG, Bydlowski SP, Chiattone CS and 
Maranhao RC (2004) Metabolism of a cholesterol-rich microemulsion (LDE) in patients 
with multiple myeloma and a preliminary clinical study of LDE as a drug vehicle for 
the treatment of the disease. Cancer Chemother Pharmacol 53:51-60. 
60. Reisinger RE and Atkinson D (1990) Phospholipid/cholesteryl ester microemulsion containing 
unesterified cholesterol: Model systems for low density lipoproteins. / Lipid Res 
31:849-858. 
186 Liu, Wang & Lu 
61. Chun PW, Brumbauge EE and Shiremann RB (1986) Interaction of human low density 
lipoprotein and apolipoprotein B with ternary lipid microemulsion. Biophys Chem 
25:223-241. 
62. Maranhao RC, Cesar TB, Pedroso-Mariani SR, Hirata MH and Mesquita CH (1993) 
Metabolic behavior in rats of a nonprotein microemulsion resembling low-density 
lipoprotein. Lipids 28:691-695. 
63. Pan G, Shawer M, 0 i e S and Lu DR (2003) In vitro gene transfection to glioma cells using 
a novel and less cytotoxic artificial lipoprotein delivery system. Pharm Res 20:738-745. 
64. Alanazi F, Fu ZF and Lu DR (2004) Effective transfection of rabies DNA vaccine in cell 
culture using an artificial lipoprotein carrier system. Pharm Res 21:676-683. 
9 
Solid Lipid Nanoparticles 
as Drug Carriers 
Karsten Mader 
1. Introduction: History and Concept of SLN 
Nanosized drug delivery systems have been developed to overcome one or several 
of the following problems: (i) low and highly variable drug concentrations 
after peroral administration due to poor absorption, rapid metabolism and elimination 
(ii) poor drug solubility which excludes i.v. injection of an aqueous drug 
solution (iii) drug distribution to other tissues combined with high toxicity (e.g. 
cancer drugs). Several systems, including micelles, liposomes, polymer nanoparticles, 
nanoemulsions and nanocapsules have been developed. During the last few 
years, solid lipid nanodispersions (SLN) have attracted increased attention. It is the 
aim of this chapter to discuss the general features of these systems with respect to 
manufacturing and performance. 
In the past, solid lipids have been mainly used for rectal and dermal applications. 
In the beginning of the 80s, Speiser and coworkers developed solid lipid 
microparticles (by spray drying)1 and "Nanopellets for peroral administration".2 
These Nanopellets were produced by dispersion of melted lipids with high speed 
mixers or ultrasound. The manufacturing process was unable to reduce all particles 
to the submicron size. A considerable amount of microparticles was present 
in the samples. This might not be a serious problem for peroral administration, 
but it excludes an intravenous injection. "Lipospheres", described by Domb, are 
187 
188 Mader 
close related systems.3-5 They are also produced by means of high shear mixing or 
ultrasound and also often contain considerable amounts of microparticles. 
The quality of the SLN has been significant improved by the use of high pressure 
homogenization (HPH) in the early 90s.6-8 Higher shear forces and a better distribution 
of the energy force more effective particle disruption, compared with high shear 
mixing and ultrasound. Dispersions obtained by this HPH are called Solid Lipid 
Nanoparticles (SLN™). Most SLN dispersions produced by high pressure homogenization 
(HPH) are characterized by an average particle size below 500 nm and 
a low microparticle content. Other production procedures are based on the use of 
organic solvents HPH/solvent evaporation9 or on dilution of microemulsions.10'11 
The ease and efficacy of manufacturing lead to an increased interest in SLN. 
Furthermore, it has been claimed that SLN combine the advantages yet without 
inheriting the disadvantages of other colloidal carriers.12,13 Proposed advantages 
include: 
• Possibility of controlled drug release and drug targeting 
• Increased drug stability 
• High drug pay load 
• Feasibility to incorporate lipophilic and hydrophilic drugs 
• No biotoxicity of the carrier 
• Avoidance of organic solvents 
• No problems with respect to large scale production and sterilization. 
However, during the last years, some of these claims have been questioned and 
it became evident that SLN are rather complex systems which possess not only 
advantages but also serious limitations. 
2. Solid Lipid Nanoparticles (SLN) Ingredients 
and Production 
2.1. General ingredien ts 
General ingredients include solid lipid(s), emulsifier(s) and water. The term lipid 
is used generally in a very broad sense and includes triglycerides (e.g. tristearine, 
hard fat), partial glycerides (e.g. Imwitor), pegylated lipids, fatty acids (stearic acid), 
steroids (e.g. cholesterol) and waxes (e.g. cetylpalmitate). All classes of emulsifiers 
(with respect to charge and molecular weight) have been used to stabilize the lipid 
dispersion. The most frequently used compounds include different kinds of poloxamer, 
polysorbates, lecithin and bile acids. It has been found that the combination 
of emulsifiers might prevent particle agglomeration more efficiently. 
Solid Lipid Nanoparticles as Drug Carriers 189 
Unfortunately, poor attention has been given by most investigators to the 
physicochemical properties of the lipid. Fatty acids, partial glycerides and other 
polar lipids are able to interact with water to much a greater extent, compared 
with a long chain triglyceride (e.g. they might form liquid crystalline phases). Polar 
lipids will have much more interaction with stabilizers (e.g. formation of mixed 
micelles), while more lipophilic lipids will show phase segregation. The author 
strongly suggests to follow the proposal by Small and to classify lipids according 
to their interactions with water.14 
2.2. SLN preparation 
2.2.1. High shear homogenization and ultrasound 
High shear homogenization and ultrasound are dispersion techniques which were 
initially used for the production of solid lipid nanodispersions.1-3 Both methods are 
widespread and easy to handle. However, dispersion quality is often poor due to the 
presence of microparticles. Furthermore, metal contamination has to be considered 
if ultrasound is used. 
Ahlin et al. used a rotor-stator homogenizer to produce SLN from different 
lipids, including trimyristin, tripalmitin, tristearin, partial glycerides 
(Witepsol®W35, Witepsol®H35) and glycerol tribehenate (Compritol®888) by meltemulsification.
15 They investigated the influence of different process parameters, 
including emulsification time, stirring rate and cooling conditions on the particle 
size and the zeta potential. Poloxamer 188 was used as steric stabilizer (0,5%w/w). 
For Witepsol®W35 dispersions, the following parameters were found to produce 
the best SLN quality: stirring 8min at 20000rpm, the optimum cooling conditions 
lOmin at 5000 rpm at room temperature. In contrary, the best conditions 
for Dynasan® 116 dispersions were 10 min emulsification at 25 000 rpm and 5 min of 
cooling at 5000 rpm in cool water (T = 16°C). An increased stirring rate did not significantly 
decrease the particle size, but improved the polydispersity index slightly. 
No general rule can be derived from differences in the established optimum emulsification 
and cooling conditions. In most cases, average particle sizes in the range 
of 100-200 nm were obtained in this study. 
2.3. High pressure homogenization (HPH) 
HPH has emerged as a very reliable and probably the most powerful technique 
for the preparation of SLN. HPH has been used for many years for the production 
of nanoemulsions for parenteral nutrition. In most cases, scaling up represents 
zero or limited problems. High pressure homogenizers push a liquid with high 
pressure (100-2000 bar) through a narrow gap (in the range of few microns). The 
190 Mader 
fluid accelerates on a very short distance to very high velocities. The high shear 
stress disrupts the particles down to the submicron range. Typical lipid contents 
are in the range of 5 to 10%. Even higher lipid concentrations (up to 40%) have been 
homogenized to lipid nanodispersions.16 
Two general approaches of the homogenization step, the hot and the cold 
homogenization techniques, can be used for the production of SLN.17,18 In both 
cases, a preparatory step involves the drug incorporation into the bulk lipid by 
dissolving the drug in the lipid melt. 
2.4. Hot homogenization 
The hot homogenization is carried out at temperatures above the melting point of 
lipid. Therefore, it is in fact the homogenization of an emulsion. A preemulsion of 
the drug loaded lipid melt and the aqueous emulsifier phase (same temperature) is 
obtained by high-shear mixing device (Ultraturrax). The quality of the preemulsion 
is very important for the final product quality. In general, higher temperatures 
result in lower particle sizes due to the decrease of the viscosity of the inner phase.19 
However, high temperatures may also increase the degradation rate of the drug and 
the carrier. The homogenization step can be repeated several times. It should be kept 
in mind however, that HPH increases the temperature of the sample (approximately 
10°C for 500 bar20). In most cases, 3 to 5 homogenization cycles at 500 to 1500 bar are 
sufficient. Increasing the homogenization pressure or the number of cycles often 
results in an increase of the particle size due to particle coalescence, which occurs 
as a result of the high kinetic energy of the particles.21 
It is important to note that the primary product of the hot homogenization is a 
nanoemulsion due to the liquid state of the lipid. Solid particles are expected to be 
formed by the following cooling of the sample to room temperature, or to temperatures 
below. Due to the small particle size and the presence of emulsifiers, lipid 
crystallization may be highly retarded and the sample may remain as a supercooled 
melt for several months.22 
2.5. Cold homogeniza Hon 
Cold homogenization has been developed to overcome the following three problems 
of the hot homogenization technique: 
(1) Temperature induced drug degradation 
(2) Drug distribution into the aqueous phase during homogenization 
(3) Complexity of the crystallization step of the nanoemulsion, leading to several 
modifications and/or supercooled melts 
The first preparatory step for cold homogenization is the same as in the hot homogenization 
procedure and includes the solubilization of the drug in the melt of the 
Solid Lipid Nanoparticles as Drug Carriers 191 
bulk lipid. However, the following steps differ. The drug containing melt is rapidly 
cooled. The high cooling rate favors a homogenous distribution of the drug within 
the lipid matrix. The solid, drug containing lipid is milled to microparticles. Typical 
particle sizes obtained by means of ball or mortar milling are in the range of 
50 to 100 microns. Low temperatures increase the fragility of the lipid, and therefore 
favor particle disruption. The solid lipid microparticles are suspended in a 
chilled emulsifier solution. The preemulsion is subjected to HPH at or below room 
temperature. An effective temperature control and regulation is needed in order to 
ensure the unmolten state of the lipid due to the increase in temperature during 
homogenization.20 In general, compared with hot homogenization, larger particle 
sizes and a broader size distribution are observed in cold homogenized samples.23 
A modified version of this technique has been recently published by the group of 
Miiller-Goymann. They dispersed a solid 1:1 lecithin-hardfat mixture (described as 
solid reversed micelles) in Tween containing water using high pressure homogenization.
24 
2.5.1. SLN prepared by solvent emulsification/evaporation 
The solvent emulsification/evaporation processes adapts techniques which have 
been previously used for the production of polymeric micro- and nanoparticles. 
The solid lipid is dissolved in a water-immiscible organic solvent (e.g. cyclohexane, 
or chloroform) that is emulsified in an aqueous phase. Upon evaporation of the solvent, 
a nanoparticle dispersion is formed by precipitation of the lipid in the aqueous 
medium. Westesen prepared nanoparticles of tripalmitate by dissolving the triglyceride 
in chloroform.25 This solution was emulsified into an aqueous phase by high 
pressure homogenization. The organic solvent was removed from the emulsion by 
evaporation under reduced pressure. The mean particle size ranges from approximately 
30 to lOOnm depending on the lecithin/co-surfactant blend. Particles with 
very small diameters (30 nm) were obtained by using bile salts as co-surfactants. 
Comparable small particle size distributions were not achievable by melt emulsification 
of similar composition. The mean particle size depends on the concentration 
of the lipid in the organic phase. Very small particles could only be obtained with 
low fat loads (5 w%) related to the organic solvent. With increasing lipid content, the 
efficacy of the homogenization declines due to the higher viscosity of the dispersed 
phase. 
2.5.2. SLN preparations by solvent injection 
The solvent injection method has been developed by Fessi to produce polymer 
nanoparticles.26 Nanoparticles were only produced with solvents which distribute 
very rapidly into the aqueous phase (e.g. ethanol, acetone, DMSO), while larger 
192 Mader 
particle sizes were obtained with more lipophilic solvents. According to Fessi, the 
particle size is critically determined by the velocity of the distribution processes and 
only water miscible solvents can be used. The solvent injection method can also be 
used for the production of solid lipid nanoparticles.27'28 However, the method is 
limited to lipids which dissolve in the polar organic solvent. Advantages of the 
method are the avoidance of elevated temperatures and high shear stress. However, 
the lipid concentration in the primary suspension will be less compared with 
High-Pressure-Homogenization. Furthermore, the use of organic solvents clearly 
represents a drawback of the method. 
2.5.3. SLN preparations by dilution of microemulsions 
or liquid crystalline phases 
SLN preparation techniques which are based on the dilution of microemulsions 
have been developed by Gasco and coworkers. Unfortunately, there is no common 
agreement within the scientific community about the definition of a microemulsion. 
One part of the scientific community understands under microemulsions 
high fluctuating systems which can be regarded as a critical solution, and therefore 
do not contain an inner and outer phase. This model has been confirmed 
by self-diffusion NMR studies of Lindman.29 In contrast, Gasco and other scientists 
understand microemulsions as two systems composed of an inner and 
outer phase (e.g. O/W-microemulsions). They are made by stirring an optical 
transparent mixture at 65-70°C, typically composed of a low melting lipid fatty 
acid (e.g. stearic acid), emulsifier (e.g. polysorbate 20, polysorbate 60, soy phosphatidylcholin, 
taurodeoxycholic acid sodium salt), co-emulsifiers (e.g. Butanol, 
Na-monooctylphosphate), and water. The hot microemulsion is dispersed in cold 
water (2-3°C) under stirring. Typical volume ratios of the hot microemulsion to 
the cold water are in the range of 1:25 to 1:50. The dilution process is critically 
determined by the composition of the microemulsion. According to the literature, 
the droplet structure is already contained in the microemulsion, and therefore, no 
energy is required to achieve submicron particle sizes.30,31 The temperature gradient 
and the pH-value determine the product quality in addition to the composition 
of the microemulsion. High temperature gradients facilitate rapid lipid crystallization 
and prevent aggregation.32'33 Due to the dilution step, lipid contents which are 
achievable are considerably lower, compared with the HPH based formulations. 
Another disadvantage includes the use of organic solvents. 
Recent work describes a similar approach to produce SLN. A hot liquid crystalline 
phase (instead of a microemulsion) is diluted in cold water to yield a solid 
lipid nanodispersion.34 This approach avoids the use of high pressure homogenization 
and organic solvents, and therefore represents an interesting opportunity. 
Solid Lipid Nanoparticles as Drug Carriers 193 
2.6. Further processing 
2.6A. Sterilization 
Sterility is required for parenteral formulations. Dry or wet heat, filtration, 
y-irradiation, chemical sterilization and aseptic production are general, opportunities 
to achieve sterility. The sterilization should not change the properties of the 
sample with respect to physical and chemical stability and the drug release kinetics. 
Sterilization by heat is a reliable procedure which is most commonly used. It was 
also applied for Liposomes.35,36 Steam sterilization will cause the formation of an 
oil in water emulsion, due to the melting of the lipid particles. The formation of SLN 
requires recrystallization of the lipids. Concerns are related to temperature induced 
changes of the physical and chemical stability. The correct choice of the emulsifier 
is of significant importance for the physical stability of the sample at high temperatures. 
Increased temperatures will affect the mobility and the hydrophilicity of 
all emulsifiers, but to a different extent. Schwarz found that Lecithin is preferable 
to Poloxamer for steam sterilization, as only a minor increase in the particle size 
and the number of microparticles was observed after steam sterilization.37'38 An 
increase in particle size for Poloxamer 188 stabilized Compritol-SLN was observed 
after steam sterilization. It was found that a decrease of the sterilization temperature 
from 121°C to 110°C can reduce sterilization induced particle aggregation to a 
large extent. This destabilization can be attributed to the decreased steric destabilization 
of the Poloxamer. It is well known for PEG-based emulsifiers that increased 
temperatures lead to dehydration of the ethylenoxide chains, pointing to a decrease 
of the thickness of the protecting layer. It has been demonstrated by 1H-NMR spectroscopy 
on Poloxamer stabilized lipid nanoparticles, that even a moderate temperature 
increase from RT to 37°C decreases the mobility of the ethylenoxide chains 
on the particle surface.39 Results of Freitas et dl. indicate that the lowering of the 
lipid content (to 2%), and the surface modification of the glass vials and nitrogen 
purging might prevent the particle growth to a large extent and avoid gelation.40 
Further studies of Cavalli et al.4* and Heiati42 demonstrate the possibility of steam 
sterilization of drug loaded SLN. 
Filtration sterilization of dispersed systems requires very high pressure and is 
not applicable to particles >0, 2 /nm. As most SLN particles are close to this size, 
filtration is of no practical use, due to the clocking of the filters. Few studies investigated 
the possibility of y-sterilization. It must be kept in mind that free radicals 
are formed during y-sterilization in all samples, due to the high energy of the yrays. 
These radicals may recombine with no modification of the sample or undergo 
secondary reactions which might lead to chemical modifications of the sample. 
The degree of sample degradation depends on the general chemical reactivity and 
the molecular mobility and the presence of oxygen. It is therefore not surprising 
194 Mader 
that chemical changes of the lipid bilayer components of liposomes were observed 
after y-irradiation.43 Schwarz investigated the impact of different sterilization techniques 
[steam sterilization at 121°C (15min) and 110°C (15min); y-sterilization] on 
SLN characteristics.37'38 In comparison to lecithin stabilized systems, Poloxamer 
stabilized SLN were less stable than steam sterilization. However, this difference 
was not detected for y-sterilized samples. Compared with steam sterilization at 
121 °C, the increase in particle size after y-irradiation was lower, but comparable to 
that at 110°C. 
Unfortunately, most investigators did not search for steam sterilization or irradiation 
induced chemical degradation. It should be kept in mind that degradation 
does not always cause increased particle sizes. In contrast, the formation of species 
like lysophosphatides or free fatty acids could even preserve small particle sizes, 
but might cause toxicological problems. Further studies with more focus on chemical 
degradation products are clearly necessary to permit valid statements of the 
possibilities of SLN sterilization. 
2.6.2. Drying by lyophilization, nitrogen purging and spray drying 
SLN are thermodynamic unstable systems, and therefore, particle growth has to be 
minimized. Furthermore, SLN ingredients and incorporated drugs are often unstable, 
hydrolyzing or oxidizing. The transformation of the aqueous SLN-suspension 
in a dry, redispersible powder is therefore often a necessary step to ensure storage 
stability of the samples. Lyophilization is widely used and is a promising way 
to increase chemical and physical SLN stability over extended periods of time. 
Lyophilization also offers principle possibilities for SLN incorporation into pellets, 
tablets or capsules. 
Two additional transformations are necessary which might be the source of 
additional stability problems. The first transformation, from aqueous dispersion to 
powder, involves the freezing of the sample and the evaporation of water under vacuum. 
Freezing of the sample might cause stability problems due to the freezing out 
effect which results in the changes of the osmolarity and the pH. The second transformation, 
resolubilization, involves situations at least in its initial stages which 
favor particle aggregation (i.e. low water and high particle content, high osmotic 
pressure). 
The protective effect of the surfactant can be compromised by lyophilization.44 
It has been found that the lipid content of the SLN dispersion should not exceed 
5%, so as to prevent an increase in the particle size. Direct contact of lipid particles 
are decreased in diluted samples. Furthermore, diluted SLN dispersions 
will also have higher sublimation velocities and a higher specific surface area.45 
The addition of cryoprotectors (e.g. Sorbitol, Mannose, Trehalose, Glucose, and 
Solid Lipid Nanoparticles as Drug Carriers 195 
Polyvinylpyrrolidon) will be necessary to decrease SLN aggregation and to obtain 
a better redispersion of the dry product. Schwarz et al. investigated the lyophilization 
of SLN in detail.46 Best results were obtained with the cryoprotectors, Glucose, 
Mannose, Maltose and Trehalose, in the concentration range between 10% and 
15%. The observations come into line with the results of the studies on liposome 
lyophilization, which indicated that Trehalose was the most sufficient substance 
to prevent liposome fusion and the leakage of the incorporated drug.47 Encouraging 
results obtained with unloaded SLN cannot predict the quality of drug loaded 
lyophilizates. Even low concentrations of 1% Tetracain or Etomidat caused a significant 
increase in particle size, excluding an intravenous administration.46 
Westesen investigated the lyophilization of tripalmitate-SLN using glucose, 
sucrose, maltose and trehalose as cryoprotective agents.48 Handshaking of redispersed 
samples was an insufficient method, but bath sonification produced better 
results. Average particle sizes of all lyophilized samples with cryoprotective agents 
were 1.5 to 2.4 times higher than the original dispersions. One year storage caused 
increased particle sizes of 4 to 6.5 times compared with the original dispersion. 
In contrast to the lyophilizates, the aqueous dispersions of tyloxapol/phospholid 
stabilized tripalmitate SLN exhibited remarkable storage stability. The instability 
of the SLN lyophilizates can be explained by the sintering of the particles. TEM pictures 
of tripalmitate SLN show an anisometrical, platelet-like shape of the particles. 
Lyophilization changes the properties of the surfactant layer due to the removal of 
water, and increases the particle concentration which favors particle aggregation. 
Increased particle sizes after lyophilization (2.1 to 4.9 times) were also reported by 
Cavalli.41 Heiati compared the influence of four cryoprotectors (i.e. trehalose, glucose, 
lactose and mannitol) on the particle size of azidothymidine palmitate loaded 
SLN lyophilizates.42 In agreement to other reports, Trehalose was found to be the 
most effective cryoprotectant. The freezing procedure will affect the crystal structure 
and the properties of the lyophilizate. Literature data suggest that the freezing 
process needs to be optimized to a particular sample size. Schwarz recommended 
rapid freezing in liquid nitrogen.46 In contrast, other researchers observed the best 
results after a slow freezing process.49 Again, best results were obtained with samples 
of low lipid content and with the cryoprotector trehalose. Slow freezing in a 
deep freeze (—70°C) was superior to rapid cooling in liquid nitrogen. Furthermore, 
introduction of an additional thermal treatment of the frozen SLN dispersion (2 hr at 
—22°C; followed by 2 hr temperature decrease to — 40°C) was found to improve the 
quality of the lyophilizate. Lately, lyophilization has been used to stabilize retinoic 
acid loaded SLN.50 
An interesting alternative to lyophilization has been recently suggested by 
Gasco's group. Drying with a nitrogen stream at low temperatures of 3 to 10°C 
has been found to be superior.51 Compared with lyophilization, the advantages of 
196 Mader 
this process are the avoidance of freezing and the energy efficiency resulting from 
the higher vapor pressure of water. 
Spray drying has been scarcely for SLN drying, although it is cheaper compared 
with lyophilization. Freitas obtained a redispersable powder with this method, 
which meets the general requirements of i.v.-injections, with regard to the particle 
size and the selection of the ingredients.52 Spray drying might potentially cause 
particle aggregation due to high temperatures, shear forces and partial melting 
of the particles. Freitas recommends the use of lipids with high melting points 
>70°C to avoid sticking and aggregation problems. Furthermore, the addition of 
carbohydrates and low lipid contents favor the preservation of the colloidal particle 
size in spray drying. 
3. SLN Structure and Characterization 
The characterization of SLN is a necessity and a great challenge. Lipid characterization 
itself is not trivial as the statement by Laggner shows53: "Lipids and fats, as soft 
condensed material in general, are very complex systems, which not only in their 
static structures but also with respect to their kinetics of supramolecular formation, 
Hysteresis phenomena or supercooling can gravely complicate the task of defining 
the underlying structures and boundaries in a phase diagram". This is especially 
true for lipids in the colloidal size range. Therefore, possible artifacts caused by sample 
preparation (removal of emulsifier from particle surface by dilution, induction 
of crystallization processes, changes of lipid modifications) should be kept in mind. 
For example, the contact of the SLN dispersion with new surfaces (e.g. a syringe 
needle) might induce lipid crystallization or modification, and sometimes result in 
the spontaneous transformation of the low viscous SLN-dispersion into a viscous 
gel. The most important parameters of SLN include particle size and shape, the 
kind of lipid modification and the degree of crystallization, and the surface charge. 
Photon correlation spectroscopy (PCS) and Laser Diffraction (LD) are the most 
powerful techniques for routine measurements of particle size. It should be kept in 
mind that both methods are not "measuring" particle sizes. Rather, they detect 
light scattering effects which are used to calculate particle sizes. For example, 
uncertainties may result from nonspherical particle shapes. Platelet structures commonly 
occur during lipid crystallization54 and are very often described in the SLN 
literature.55-59 The influence of the particle shape on the measured size is discussed 
by Sjostrom.55 Further difficulties arise both in PCS and LD measurements for samples 
which contain several populations of different size. Therefore, additional techniques 
might be useful. For example, light microscopy is recommended although 
it is not sensitive to the nanometer size range. It gives a fast indication about the 
Solid Lipid Nanoparticles as Drug Carriers 197 
presence and the character of microparticles. Electron Microscopy provides, in contrast 
to PCS and LD, direct information on the particle shape.57'58 Atomic force 
microscopy (AFM) has attracted increasing attention. A cautionary note applies to 
the use of AFM in the field of nanoparticles, as an immobilization of the SLN by 
solvent removal is required to assess their shape by the AFM tip. This procedure is 
likely to cause substantial changes of the molecular structure of the particles. Zur 
Miihlen demonstrated the ability of AFM to image the morphological structure of 
SLN.60 The sizes of the visualized particles are of the same magnitude, compared 
with the results of PCS measurements. The AFM investigations revealed the disklike 
structure of the particles. Dingier investigated cetylpalmitate SLN (stabilized 
by polyglycerol methylglucose distearate, Tego Care 450) by electron microscopy 
and AFM and found an almost spherical form of the particles.61 The usefulness of 
cross flow Field-Flow-Fractionation (FFF) for the characterization of colloidal lipid 
nanodispersions has been recently demonstrated.58 Lipid nanodispersions with 
constant lipid content, but different ratios of liquid and solid lipids did show similar 
particle sizes in dynamic light scattering. However, retention times in FFF were 
remarkably dissimilar due to the different particle shapes (i.e. spheres vs. platelets). 
Anisotropic particles such as platelets will be constrained by the cross flow much 
more heavily compared with the spheres of similar size. The very high anisometry 
of the SLN particles has been confirmed by electron microscopy, where very thin 
particles of 15 nm thickness and the length of several hundred nanometers became 
visible. 
The measurement of the zeta potential allows predictions about the storage 
stability of colloidal dispersions.62 In general, particle aggregation is less likely 
to occur for charged particles (i.e. high zeta potential) due to electric repulsion. 
However, this rule cannot strictly apply to systems which contain steric stabilizers, 
because the adsorption of steric stabilizer will decrease the zeta potential due to the 
shift in the shear plane of the particle. 
Particle size analysis is just one aspect of SLN quality. The same attention has to 
be paid on the characterization of lipid crystallinity and modification, because these 
parameters are strongly correlated with drug incorporation and release rates. Thermodynamic 
stability and lipid packing density increase, and drug incorporation 
rates decrease in the following order: 
supercooled melt < a-modification < B'-modification < 6-modification 
In general, it has been found that melting and crystallization processes of 
nanoscaled material can differ considerable from that of the bulk material.63 The 
thermodynamic properties of material having small nanometer dimensions can be 
considerably different, compared with the material in bulk form (e.g. the reduction 
198 Mader 
of melting point). This occurs because of the tremendous influence of the surface 
energy. 
This statement is also valid for SLN, where lipid crystallization and modification 
changes might be highly retarded,64 due to the small size of the particles and 
the presence of emulsifiers. Moreover, crystallization might not occur at all and 
has been shown that samples which were previously described as SLN (solid lipid 
particles) were in fact supercooled melts (liquid lipid droplets).65 The impact of 
the emulsifier on SLN lipid crystallization has been shown by Bunjes.66 The same 
group demonstrated also a size dependent melting of SLN.67 
Differential Scanning Calorimetry (DSC) and X-ray scattering are most commonly 
applied to asses the status of the lipid. DSC uses the fact that different lipid 
modifications possess different melting points and melting enthalpies. By means 
of X-ray scattering, it is possible to assess the length of the long and short spacings 
of the lipid lattice. It is highly recommended to measure the SLN dispersion 
themselves, because solvent removal will lead to modification changes. Sensitivity 
problems and long measurement times of convential X-ray sources might be 
overcome by synchrotron irradiation.64 In addition, this method permits to conduct 
time resolved experiments and allows the detection of intermediate states 
of colloidal systems which will be non detectable by convential X-ray methods.53 
Recent work shows that SLN might form superstructures by parallel alignment of 
SLN platelets. These reversible particle self-assemblies were observed by Illing 
et al. in tripalmitin dispersions when the lipid concentration exceeds 40mg/g. 
Higher lipid concentrations did enhance particle self-assembly. The tendency to 
form self-assemblies has been found to depend on the particle shape, the lipid 
and the surfactant concentration.68 Infrared and Raman Spectroscopy are useful 
tools to investigate structural properties of lipids and they might give complentary 
information to X-ray and DSC.54 Raman measurements on SLN show that the 
arrangement of lipid chains of SLN dispersions changes with storage.69 
Rheometry might be particularly useful for the characterization of the viscoelastic 
properties of SLN dispersions. The rheological properties are important with 
respect to the dermatological use of SLN, but they also provide useful information 
about the structural features of SLN dispersions and their storage dependency. 
Studies of Lippacher show that the SLN dispersion posses higher elastic properties 
than emulsions of comparable lipid content.70-72 Furthermore, a sharp increase of 
the elastic module is observed at a certain lipid content. This point indicates the 
transformation from a low viscous lipid dispersion to an elastic system, with a 
continuous network of lipid nanocrystals. Illing and Unruh did compare the rheological 
properties of trimyristic, tripalmitic and tristearic SLN suspensions. The 
results indicate that the viscosity of triglyceride suspensions increases with the 
lipid chain length and an increased anisotropy of the particles.73 Souto et al. used 
Solid Lipid Nanoparticles as Drug Carriers 199 
rheology to study the influence of SLN addition on the rheological properties of 
hydrogels.74 
The co-existence of additional colloidal structures (micelles, liposomes, mixed 
micelles, nanodispersed liquid crystalline phases, supercooled melts, drugnanoparticles) 
has to be taken into account for all SLN dispersions. Unfortunately, 
many investigators neglect this aspect, although the total amount of surface active 
compounds is often comparable to the total amount of the lipid. The characterization 
and quantification are serious challenges due to the similarities in size. In addition, 
the sample preparation will modify the equilibrium of the complex colloidal 
system. Dilution of the original SLN dispersion with water might cause the removal 
of surfactant molecules from the particle surface and induce further changes such 
as crystallization or the changes of the lipid modifications. It is therefore highly 
desirable to use methods which are sensitive to the simultaneous detection of different 
colloidal species, which do not require preparatory steps such as Raman, 
NMR and ESR spectroscopy. 
NMR active nuclei of interest are 1H, 13C, 19F and 35P. Due to the different chemical 
shifts, it is possible to attribute the NMR signals to particular molecules or their 
segments. For example, lipid methyl protons give signals at 0.9 ppm, while protons 
of the polyethylenglycole chains give signals at 3.7 ppm. Simple ^-spectroscopy 
permits an easy and rapid detection of supercooled melts, due to the low linewidths 
of the lipid protons69,75-77. This method is based on the different proton relaxation 
times in the liquid and semisolid/solid state. Protons in the liquid state give sharp 
signals with high signal amplitudes, while semisolid/solid protons give very broad 
or invisible NMR signals under these circumstances. NMR has been used to characterize 
calixarene SLN78 and hybrid lipid particles (NLC), which are composed 
of liquid and solid lipids.59 Protons from solid lipids are not detected by standard 
NMR, but they can be visualized by solid state NMR. A drawback of solid 
state NMR is the rapid spinning of the sample that might cause artifacts. A recent 
paper describes the use of this method to monitor the distribution of Q10 in lipid 
matrices.79 Unfortunately, the authors did use "drying of the sample to constant 
weight" as a preparatory step, which will cause significant changes of the sample 
characteristics. 
ESR requires the addition of paramagnetic spin probes to investigate SLN dispersions. 
A large variety of spin probes is commercially available. The corresponding 
ESR spectra give information about the microviscosity and micropolarity. ESR 
permits the direct, repeatable and non-invasive characterization of the distribution 
of the spin probe between the aqueous and the lipid phase.80 Experimental results 
demonstrate that storage induced crystallization of SLN leads to an expulsion of 
the probe out of the lipid into the aqueous phase.81 Furthermore, using an ascorbic 
acid reduction assay, it is possible to monitor the time scale of the exchange between 
200 Mader 
the aqueous and the lipid phase.59 The transfer rates of molecules between SLN and 
liposomes or cells have been determined by ESR.82 
4. The "Frozen Emulsion Model" and Alternative SLN Models 
Lipid nanoemulsions are composed of a liquid oily core and a surfactant layer 
(lecithin). They are widely used for the parenteral delivery of poorly soluble 
drugs.83-85 The original idea of SLN was to achieve a controlled release of incorporated 
drugs by increasing the viscosity of the lipid matrix. Therefore it is not 
surprising that in original model, SLN is being described as "frozen emulsions" 
(see Fig. 1, left and middle).8687 However, lipids are known to crystallize very frequently 
in anisotropic platelet shapes54 and anisotropic. Sjostrom et al. described 
in 1995 that the particle shape of Cholesterylacetate SLN did strongly depend on 
the emulsifier.55 Platelet shaped particles have been detected for lecithin stabilized 
particles, while PEG-20-sorbitanmonolaurate stabilized particles preserved their 
spherical shape. Anisotropic particles have been found in numerous other SLN 
dispersions.56-59 Based on the experimental results, a platelet shaped SLN model 
can be proposed as an alternative (see Fig. 1, right). 
In the year 2000, Westesen questioned the frozen emulsion droplet model with 
the following statement88: 
"Careful physicochemical characterization has demonstrated that these lipid-based 
nanosuspensions (solid lipid nanoparticles) are not just emulsions with solidified 
droplets. 
During the development process of these systems, interesting phenomena have 
been observed, such as gel formation on solidification and upon storage, unexpected 
dynamics of polymorphic transitions, extensive annealing of nanocrystals 
over significant periods of time, stepwise melting of particle fractions in the 
Nanoemulsion SLN: "Frozen emulsion droplet" SLN: Platelet shaped particles o o — 
Core: liquid lipid (oil) S Core: solid lipid H Shell: stabilizer 
Fig. 1. General structure of a nanoemulsion (left), and proposed models for SLN: Frozen 
emulsion droplet model (middle) and platelet shaped SLN model (right). 
Solid Lipid Nanoparticles as Drug Carriers 201 
lower-nanometer-size range, drug expulsion from the carrier particles on crystallization 
and upon storage, and extensive supercooling." 
Her comment highlights the complex behavior and changes of SLN dispersions. 
In addition, the presence of competing colloidal structures (e.g. micelles, 
liposomes, mixed micelles, nanodispersed liquid crystalline phases, supercooled 
melts and drug-nanoparticles) should be considered. Additional colloids might 
have an impact on very different aspects, including the correct measurement of 
particle size, drug incorporation and toxicity. A recent study shows that the cell 
toxicity of the SLN dispersion was reduced by dialysis due to the removal of water 
soluble components.89 
5. Nanostructured Lipid Carriers (NLC) 
Nanostructured lipid carriers (NLC) have been recently proposed as a new SLN 
generation with improved characteristics.90 The general idea behind the system is 
to improve the poor drug loading capacity of SLN by "mixing solid lipids with 
spatially incompatible lipids leading to special structures of the lipid matrix",91 
while still preserving controlled release features of the particles. Three different 
types of NLC have been proposed (NLC I: The imperfect structured type, NLC 
II: The structureless type and NLC III: The multiple type). Unfortunately, these 
structural proposals have not been supported by experimental data. They assume 
a spherical shape and they are not compatible with lipid platelet structures. 
For example, NLC III structures should contain small oily droplets in a solid 
lipid sphere (Fig. 2, left). Detailed analytical examination of NLC systems by Jores 
et al. demonstrate that "nanospoon" structures are formed, in which the liquid oil 
adheres on the solid surface of a lipid platelet (Fig. 2, right). 
Jores et al. did conclude that "Neither SLN nor NLC lipid nanoparticles showed 
any advantage with respect to incorporation rate or retarded accessibility to the 
drug, compared with conventional nanoemulsions. The experimental data concludes 
that NLCs are not spherical solid lipid particles with embedded liquid 
liquid lipid (oil) J A solid lipid • stabilizer 
Fig. 2. Proposed NLC III structure (modified after91) and experimental determined 
"nanospoon" structure described by Jores et al. (side view of particle).58'59 
202 Mader 
droplets, but rather, they are solid platelets with oil present between the solid 
platelet and the surfactant layer". Very similar structures have been found on Q10 
loaded SLN by Bunjes et til.92 
6. Drug Localization and Release 
Proposed advantages of SLN, compared with nanoemulsions, include increased 
protection capacity against drug degradation and controlled release possibilities 
due to the solid lipid matrix. The general low capacity of crystalline structures to 
accommodate foreign molecules is a strong argument against the proposed rewards. 
It is therefore necessary to distinguish between drug association and drug incorporation. 
Drug association means that the drug is associated with the lipid, but it 
might be localized in the surfactant layer or between the solid lipid and the surfactant 
layer (similar to the oil in Fig. 2, right). Drug incorporation would mean 
the distribution of the drug within the lipid matrix. Another limiting aspect comes 
from the fact that the platelet structure of SLN, which is found in many systems, 
leads to a tremendous increase in surface area and the shortening of the diffusion 
lengths. Furthermore, additional colloid structures present in the sample are 
alternatives for drug localization the SLN for drug incorporation as it was pointed 
out by Westesen88: "The estimation of drug distribution is difficult for dispersions 
consisting of more than one type of colloidal particle. Depending on the type of 
stabilizer and on the concentration ratio of stabilizer to matrix material significant 
numbers of particles such as liposomes and/or (mixed) micelles may coexist with 
the expected type of particles". 
The detailed investigation of drug localization is very difficult and only a few 
studies exist. Parelectric spectroscopy has been used to investigate the localization 
of glucocorticoids. The results indicate that the drug molecules are attached to the 
particle surface, but not incorporated into the lipid matrix. With Betamethasonvalerate, 
the loading capacity of the particle surface was clearly below the usual concentration 
of 0.1%.93 Lukowski used Energy Dispersive X-ray Analysis and found 
that the drug Triamcinolone, Dexamethasone and Chloramphenicol are partially 
stored at the surface of the individual nanoparticles.94 
The importance of the emulsifier is reflected in a study from Danish scientists.95 
They produced gamma-cyhalothrin (GCH) loaded lipid micro- and nanoparticles. 
GCH had only limited solubility in the solid lipid and was expulsed during storage. 
The appearance of GCH crystals was strongly dependent from the solubility 
of the GCH in the emulsifier solutions. Emulsifier with high GCH solubility provoked 
rapid crystal growth. This observation is in accordance with a mechanism of 
crystal growth according to Ostwald ripening. Slovenian scientist found that ascorbylpalmitate 
was more resistant against oxidation in non-hydrogenated soybean 
Solid Lipid Nanoparticles as Drug Carriers 203 
lecithin liposomes, compared with SLN.96 It shows that liposomes might have a 
higher protection capacity compared with SLN. 
Fluorescence and ESR studies have been used by Jores et al. to monitor the 
microenvironment and the mobility of model drugs. The results indicate that even 
highly lipophilic compounds are pushed into a polar environment during lipid 
crystallization. Therefore, the incorporation capacity of SLN is very poor for most 
molecules.69 A nitroxide reduction assay gave results in accordance with the results 
of the distribution. Compared with nanoemulsions, nitroxides were more accessible 
in SLN and NLC to ascorbic acid, localized in the aqueous environment. Therefore, 
nanoemulsions were more protective than SLN and NLC systems. 
Drug release from SLN and NLC could be either controlled by the diffusion of 
the drug or the erosion of the matrix. The original idea was to achieve a controlled 
release of SLN due to the slowing down of drug diffusion to the particle surface. This 
idea is, however, questionable due to drug expulsion during lipid crystallization. 
In addition, very short diffusion lengths in nanoscaled delivery systems lead to 
short diffusion times, even in highly viscous or solid matrices. In most cases, the 
delivery of the drug will be controlled by the slow dissolution rate in the aqueous 
environment. Drug release rate will be highly dependent on the presence of further 
solubilizing colloids (e.g. micelles), which are able to work as a shuttle for the drug 
and the presence or absence of a suitable acceptor compartment. Many investigators 
studied only the release in buffer media. A controlled release pattern under such 
conditions is not surprising, as it is caused by low solubilization kinetics due to 
the poor solubility of the drug. In vivo, acceptor compartments will be present 
(e.g. lipoproteins, membranes) and will speed up release processes significantly. 
Whenever possible, drug loaded SLN should be compared with nanosuspensions 
to separate the general features of the drug and the influence of the lipid matrix. 
Results by Kristl et al. indicate that lipophilic nitroxides diffuse between SLN 
and liposomes. The diffusion kinetics was strongly dependent on the nitroxide 
structure. In contrast, uptake of nitroxides in cells was similar between lipophilic 
nitroxides, suggesting endocytosis as the main mechanism.82 The detailed mechanisms 
of drug release in vivo are poorly understood. In vitro data by Olbrich demonstrate 
that SLN are degraded by lipases.97,98 Degradation by lipase depends on the 
lipid and strongly on the surfactant. Steric stabilization (e.g. by poloxamer) of SLN 
and NLC are less accessible because lipase needs an interface for activation. It is 
also known that highly crystalline lipids are poorly degraded by lipase. 
7. Administration Routes and In Vivo Data 
SLN and NLC can be administrated at different routes, including peroral, dermal, 
intravenously and pulmonal. Peroral administration of SLN could enhance the drug 
204 Mader 
absorption and modify the absorption kinetics. Despite the fact that in most of the 
SLN, the drug will be associated but not incorporated in the lipid, SLN might have 
advantages due to enhanced lymphatic uptake, enhanced bioadhesion or increased 
drug solubilization by SLN lipolysis products such as fatty acids and monoglycerides. 
A serious challenge represents the preservation of the colloidal particle in 
the stomach, where low pH values and high ionic strengths favor agglomeration 
and particle growth. Zimmermann and Muller studied the stability of different 
SLN formulations in artificial gastric juice." The main findings of this study are 
that (i) some SLN dispersions preserve their particle size under acidic conditions, 
and (ii) there is no general lipid and surfactant which are superior to others. The 
particular interactions between lipid and stabilizer are determining the robustness 
of the formulation. Therefore, the suitable combination of ingredients has to be 
determined on a case by case basis. 
Several animal studies show increased absorption of poorly soluble drugs. The 
efficacy of orally administrated Triptolide free drug and Triptolide loaded SLN 
have compared in the carrageenan-induced rat paw edema by Mei et al.wo Their 
results suggest that SLN can enhance the anti inflammatory activity of triptolide 
and decrease triptolide-induced hepatotoxicity. The usefulness of SLN to increase 
the absorption of the poorly soluble drug all-trans retinoic acid has been shown 
by Hu et al. on rats.101 Gascos group investigated the uptake and distribution of 
Tobramycin loaded SLN in rats.102'103 They observed an increased uptake into the 
lymph, which causes prolonged drug residence times in the body of the animals. 
Furthermore, AUC and clearance rates did depend on the drug load. The same 
group described also enhanced absorption of Idarubicin-loaded solid lipid nanoparticles 
(IDA-SLN), in comparison to the drug solution. Furthermore, the authors 
described that SLN were able to pass the blood-brain barrier and concluded that 
duodenal administration of IDA-SLN modifies the pharmacokinetics and tissue 
distribution of idarubicin.104 
Parenteral administration of SLN is of great interest too. To avoid the rapid 
uptake of the SLN by the RES system, stealth SLN particles have been developed 
by the adoption of the stealth concept from liposomes and polymer nanoparticles. 
Reports indicate that Doxorubicin loaded stealth SLN circulate for long period of 
time in the blood and change the tissue distribution.105 Therefore, SLN could be 
alternatives to marketed stealth-liposomes, which can decrease the heart toxicity 
of this drug due to changed biodistribution. Long circulation times have also been 
observed for Poloxamer stabilized SLN with Paclitaxel.106 
The dermal application is of particular interest and it might become the main 
application of SLN.107 SLN pose occlusive properties which are related to the solid 
structure of the lipid.108 Human in vivo results of the group of Muller demonstrate 
that SLN can improve skin hydration and viscoelasticity.109 SLN have also 
Solid Lipid Nanoparticles as Drug Carriers 205 
UV protection capacity due to their reflection of UV light.110 Furthermore, data by 
Schafer-Korting suggest SLN can be used to decrease drug side effects due to SLN 
mediated drug targeting to particular skin layers.111 
Further reports describe additional applications of SLN as well as gene 
delivery,112 delivery to the eye,113 pulmonary delivery,114 and drug targeting of 
anticancer drugs.115 Studies of the different groups also propose the use of SLN for 
brain targeting to deliver MRI contrast agents116 or antitumour drugs.117'118 
8. Summary and Outlook 
SLN and NLC are now investigated by many scientists worldwide. In contradiction 
to early proposals, they certainly do not combine all the advantages of the other 
colloidal drug carriers and avoid the disadvantages of them. SLN are complex colloidal 
dispersions, not just "frozen emulsions". SLN dispersions are very susceptible 
to the sample history and storage conditions. Disadvantages of SLN include 
gel formation on solidification and upon storage, unexpected dynamics of polymorphic 
transitions, extensive annealing of nanocrystals over significant periods 
of time, stepwise melting of particle fractions in the lower-nanometer-size range, 
drug expulsion from the carrier particles on crystallization and upon storage, and 
extensive supercooling. The anisotropic shape of many SLN dispersions increases 
the surface area significantly, decreases the diffusion lengths to the surface and 
changes the rheological behavior dramatically (e.g. gel formation). Furthermore, 
the presence of alternative colloidal structures (micelles, liposomes) has to be considered 
to contribute to drug localization. In most cases, the drug will be associated 
with the lipid and not incorporated. Studies demonstrate that SLN and NLC might 
have no advantages compared with submicron emulsions, in regard to protection 
from the aqueous environment. 
On the other side, animal data suggest that SLN can change the pharmacokinetics 
and the toxicity of drugs. In many cases, drug incorporation might not be 
required and drug association with the lipid can be sufficient for lymphatic uptake. 
Clearly, more detailed studies are necessary to get a deeper understanding of the 
in vivo fate of these carriers. Whenever possible, SLN and NLC systems should 
be compared directly with alternative colloidal carriers (e.g. liposomes, nanoemulsions, 
nanosuspensions) to evaluate their true potential. 
References 
1. Eldem T, Speiser P and Hincal A (1991) Optimization of spray-dried and congealed lipid 
micropellets and characterization of their surface morphology by scanning electron 
microscopy. Pharm Res 8:47-54. 
206 Mader 
2. Speiser P (1990) Lipidnanopellets als Tragersystem fur Arzneimittel zur peroralen 
Anwendung, European Patent EP 0167825. 
3. Domb AJ (1993) Lipospheres for controlled delivery of substances. United States Patent 
No. 5188837. 
4. Domb AJ (1995) Long acting injectable oxytetracycline-liposphere formulation. Int } 
Pharm 124:271-278. 
5. Domb AJ (1993) Liposphere parenteral delivery system. Proc Intl Symp Control Rel Bioact 
Mater 20:346-347. 
6. Siekmann B and Westesen K (1992) Submicron-sized parenetral carrier systems based 
on solid lipids, Pharm. Pharmacol Lett 1:123-126. 
7. Miiller RH, Lucks JS (1996) Arzneistofftrager aus festen Lipidteilchen, Feste Lipidnanospharen 
(SLN). European Patent No. 0605497. 
8. Miiller RH, Mehnert W, Lucks JS, Schwarz C, zur Miihlen A, Weyhers H, Freitas C 
and Riihl D (1995) Solid Lipid Nanoparticles (SLN) —An Alternative Colloidal Carrier 
System for Controolled Drug Delivery. Eur J Pharm Biopharm 41:62-69. 
9. Sjostrom B and Bergenstahl B (1992) Preparation of submicron drug particles in lecithinstabilized 
o /w emulsions. I. Model studies of the precipitation of cholesteryl acetate. 
Int} Pharm 88:53-62. 
10. Cavalli R, Caputo O and Gasco MR (1993) Solid lipospheres of doxorubicin and idarubicin. 
Int]Pharm 89:R9-R12. 
11. Gasco MR (1993) Method for producing solid lipid microspheres having a narrow size 
distribution. United States Patent No. 5250236. 
12. Miiller RH (1997) Pharmazeutische Technologie, Moderne Arzneiformen, Wiss. 
Verlagsges. Stuttgart. 
13. Miiller RH and Runge SA, Solid Lipid Nanoparticles (SLN) for controlled drug delivery, 
in Submicron Emulsions in Drug Targeting and Delivery, Benita S (ed.), Harwood Academic 
Publishers. 
14. Small D (1986) Handbook of lipids. The physical chemistry of lipids: From alkanes to 
phospholipids, Plenum Press, New York. 
15. Ahlin P, Kristl J and Smid-Kobar J (1998) Optimization of procedure parameters and 
physical stability of solid lipid nanoparticles in dispersions. Acta Pharm 48:257-267. 
16. Lippacher A, Miiller RH and Mader K (2000) Investigation on the viscoelastic properties 
of lipid based colloidal drug carriers. Int ] Pharm 196:227-230. 
17. zur Miihlen A and Mehnert W (1998) Drug release and release mechanism of prednisolone 
loaded solid lipid nanoparticles. Pharmazie 53:552-555. 
18. zur Miihlen A, Schwarz C and Mehnert W (1998) Solid lipid nanoparticles (SLN) for 
controlled drug delivery — Drug release and release mechanism. 
19. Lander R, Manger W, Scouloudis M, Ku A, Davis C and Lee A (2000) Gaulin homogenization: 
A mechanistic study. Biotechnol Prog 16:80-85. 
20. Jahnke S (1998) The theory of high pressure homogenization, in Emulsions and Nanosuspensions 
for the Formulation of Poorly Soluble Drugs, Miiller RH, Benita S and Bohm B 
(eds.), Medpharm Scientific Publishers: Stuttgart, pp. 177-200. 
Solid Lipid Nanoparticles as Drug Carriers 207 
21. Siekmann B and Westesen K (1994) Melt-homogenized solid lipid nanoparticles stabilized 
by the nonionic surfactant tyloxapol, I. Preparation and particle size determination. 
Pharm Pharmacol Lett 3:194-197. 
22. Bunjes H, Siekmann B and Westesen K (1998), Emulsions of supercooled melts — a 
novel drug delivery system, in Submicron Emulsions in Drug Targeting and Delivery, 
Benita S (ed.), Harwood Academic Publishers. 
23. zur Miihlen A (1996) Feste Lipid-Nanopartikel mit prolongierter Wirkstoffliberation: 
Herstellung, Langzeitstabilitat, Charakterisierung, Freisetzungsverhalten und - 
machanismen, PhD thesis, Free University of Berlin. 
24. Friedrich I and Miiller-Goymann CC (2003) Characterization of solidified reverse micellar 
solutions (SRMS) and production development of SRMS-based nanosuspensions. 
Eur J Pharm Biopharm 56:111-119. 
25. Siekmann B and Westesen K (1996) Investigations on solid lipid nanoparticles prepared 
by precipitation in o /w emulsions. Eur } Pharm Biopharm 43:104-109. 
26. Fessi H, Puisieux F, Ammoury N and Benita S (1989) Nanocapsule formation by interfacial 
polymer deposition following solvent displacement. Int J Pharm 55:R1-R4. 
27. Hu FQ, Yuan H, Zhang HH and Fang M (2002) Preparation of solid lipid nanoparticles 
with clobetasol propionate by a novel solvent diffusion method in aqueous system and 
physicochemical characterization. Int J Pharm 239:121-128. 
28. Schubert MA and Miiller-Goymann CC (2003) Solvent injection as a new approach for 
manufacturing lipid nanoparticles—evaluation of the method and process parameters. 
Eur } Pharm Biopharm 55:125-131. 
29. Danielsson I and Lindman B (1981) The definition of microemulsion. Coll Surf B 3: 
391-392. 
30. Gasco MR (1997) Solid lipid nanospheres from warm micro-emulsions. Pharma Technol 
Eur 52-58. 
31. Boltri L, Canal T, Esposito PA and Carli F (1993) Lipid nanoparticles: Evaluation of some 
critical formulation parameters. Proc Intl Symp Control Rel Bioact Mater 20:346-347. 
32. Cavalli R, Marengo E, Rodriguez L and Gasco MR (1996) Effect of some experimental 
factors on the production process of solid lipid nanoparticles. Eur J Pharm Biopharm 
43:110-115. 
33. Gasco MR, Morel S and Carpigno R (1992) Optimization of the incorporation of desoxycortisone 
acetate in lipospheres. Eur } Pharm Biopharm 38:7-10. 
34. Dahms G and Seidel H (2004) Method for the preparation of solid-lipid nanoparticles 
(SLNs) without high pressure homogenizer for pharmaceutical, cosmetic and food 
applications. German Patant application DE 2003-10312763 20030321. 
35. Zuidam NJ, Lee SS, L and Crommelin DJA (1992) Sterilization of liposomes by heat 
treatment. Pharm Res 10:1591-1596. 
36. Lukyanov AN and Torchilin VP (1994) Autoclaving of liposomes. / Microencap 11: 
669-672. 
37. Schwarz C and Mehnert W (1995) Sterilization of drug-free and tetracaine-loaded solid 
lipid nanoparticles (SLN). Proc 1st World Meeting APGI/APV, Budapest, 485^86. 
208 Mader 
38. Schwarz C, Freitas C, Mehnert W and Muller RH (1995) Sterilization and physical 
stability of drug-free and etomidate-loaded solid lipid nanoparticles. Proc Intl Symp 
Control Rel Bioct Mater 22:766-767. 
39. Liedtke S, Jores K, Mehnert W and Mader K (2000) Possibilities of non-invasive physicochemical 
characterisation of colloidal drug carriers, 27th Intl Symp Control Rel Bioact 
Mater Vol. 27, Controlled Release Society, Paris, 1088-1089. 
40. Freitas C (1998) Feste Lipid-Nanopartikel (SLN): Mechanismen der physikalischen 
Destabilisierung und Stabilisierung. PhD thesis, Free University of Berlin. 
41. Cavalli R, Caputo O, Carlotti ME, Trotta M, Scarnecchia and Gasco MR (1997) Sterilization 
and freeze-drying of drug-free and drug-loaded solid lipid nanoparticles. Int} 
Pharm 148:47-54. 
42. Heiati H, Tawashi R and Phillips NC (1998) Drug retention and stability of solid lipid 
nanoparticles containuing azidothymidine palmitate after autoclaving, storage and 
lyophilisation.} Microencap 15:173-184. 
43. Sculier JP, Coune A, Brassine C, Laduron C, Atassi G, Ruysschert GM and 
Fruhling J (1986) Intravenous infusion of high doses of liposomes containing NSC 
251635, a water insoluble cytostatic agent. A pilot sudy with pharmacokinetic data. 
/ Clin Oncol 4:789-797. 
44. Rupprecht H (1993) Physikalisch-chemische Grundlagen der Gefriertrocknung, in 
Essig D and Oschmann R (eds.), Lyophilization. Paperback APV, Band 35, Wissenschaftliche 
Verlagsgesellschaft mbH, Stuttgart, pp. 13-38. 
45. Pikal MJ, Shah S, Roy ML and Putman R (1990) The secondary drying stage of freeze 
drying: drying kinetics as a function of temperature and chamber pressure. Int} Pharm 
60:203-217. 
46. Schwarz C and Mehnert W (1997) Freeze-drying of drug-free and drug-loaded solid 
lipid nanoparicles. Int J Pharm 157:171-179. 
47. Crowe LM, Crowe JH, Rudolph A, Womersley C and Appel L (1985) Preservation of 
Freeze-dried Liposomes by Trehalose. Arch Biochem Biophys 242:240-247. 
48. Siekmann B and Westesen K (1994) Melt-homogenized solid lipid nanoparticles stabilized 
by the nonionic surfactant tyloxapol, II. Physicochemical characterization and 
lyophilisation. Pharm Pharmacol Lett 3:225-228. 
49. Zimmermann E, Muller RH and Mader K (2000) Influence of different parameters on 
reconstitution of lyophilized SLN. Int J Pharm 196:211-213. 
50. Lim SJ, Lee MK and Kim CK (2004) Altered chemical and biological activities of all-trans 
retinoic acid incorporated in solid lipid nanoparticle powders. / Control Rel 100:53-61. 
51. Marengo E, Cavalli R, Rovero G and Gasco MR (2003) Scale-up and optimization of an 
evaporative drying process applied to aqueous dispersions of solid lipid nanoparticles. 
Pharm Dev Techn 8:299-309. 
52. Freitas C and Muller RH (1998) Spray-drying of solid lipid nanoparticles (SLN™). Eur 
J Pharm Biopharm 46:145-151. 
53. Laggner P (1999) X-ray diffraction of lipids, in Spectral Properties of Lipids, Hamilton RJ 
and Cast J (eds.), Sheffield Academic Press. 
Solid Lipid Nanoparticles as Drug Carriers 209 
54. Garti N and Sato K (eds.), (1998) Crystallization and Polymorphism of Fats and Fatty 
Acids, Marcel Dekker; New York and Basel. 
55. Sjostrom B, Kaplun A, Talmon Y and Cabane B (1995) Structures of nanoparticles prepared 
from oil in water emulsions. Pharm Res 12:39^8. 
56. Siekmann B and Westesen K (1992) Sub-micron sized parenteral carrier systems based 
on solid lipid, Pharma Pharmacol Lett 1:123-126. 
57. Illing A, Unruh T and Koch MHJ (2004) Investigation on particle self-assembly in solid 
lipid-based colloidal drug carrier systems. Pharm Res 21:592-597. 
58. Jores K, Mehnert W, Drechsler M, Bunjes H, Johann C and Mader K (2004) Investigations 
on the structure of solid lipid nanoparticles (SLN) and oil-loaded solid lipid 
nanoparticles by photon correlation spectroscopy, field-flow fractionation and transmission 
electron microscopy. / Control Rel 95:217-227. 
59. Jores K, Mehnert W and Mader K (2003) Physicochemical investigations on solid lipid 
nanoparticles (SLN) and on oil-loaded solid lipid nanoparticles: A NMR- and ESRstudy. 
Pharm Res 20:1274-1283. 
60. zur Muhlen A, zur Miihlen E, Niehus H and Mehnert W (1996) Atomic force microscopy 
studies of solid lipid nanoparticles. Pharm Res 13:1411-1416. 
61. Dingier A, Blum RP, Niehus H, Muller RH and Gohla S (1999) Solid lipid nanoparticles 
(SLN™/Lipopearls™) — a pharmaceutical and cosmetic carrier for the application of 
vitamin E in dermal products.} Microencap 16:751-767. 
62. Muller RH (1996) Zetapotential und Partikelladung - Kurze Theorie, praktische Mefidurchfuhrung, 
Dateninterpretation, Wissenschaftliche Verlagsgesellschaft Stuttgart. 
63. Lai SL, Guo JY, Petrova V, Ramanath G and Allen LH (1996) Size-Dependent Melting 
Properties of Small Tin Particles: Nanocalorimetric Measurements. Phys Rev Lett 
77:99-103. 
64. Westesen K, Siekmann B and Koch MHJ (1993) Investigations on the physical state of 
lipid nanoparticles by synchrotron radiation X-ray diffraction. Int} Pharm 93:189-199. 
65. Westesen K and Bunjes H (1995) Do nanoparticles prepared from lipids solid at room 
temperature always possess a solid matrix? Int ] Pharm 115:129-131. 
66. Bunjes H, Koch MHJ and Westesen K (2003) Influence of emulsifiers on the crystallization 
of solid lipid nanoparticles. / Pharm Sci 92:1509-1520. 
67. Bunjes H, Koch MHJ and Westesen K (2000) Effect of particle size on colloidal solid 
triglycerides. Langmuir 16:5234-5241. 
68. Illing A, Unruh T and Koch MHJ (2004) Investigation on Particle Self-Assembly in Solid 
Lipid-Based Colloidal Drug Carrier Systems. Pharm Res 21:592-597. 
69. Jores K (2004) Lipid nanodispersions as drug carrier systems — a physicochemical characterization, 
Thesis, University of Halle (http://sundoc.bibliothek.uni-halle.de/dissonline 
/ 04 / 04H310 / prom.pdf). 
70. Lippacher A, Muller RH and Mader K (2004) Liquid and semisolid SLN dispersions 
for topical application: Rheological characterization. Eur J Pharm Biopharm 58:561-567. 
71. Lippacher A, Muller RH and Mader K (2001) Preparation of semisolid drug carriers for 
topical application based on solid lipid nanoparticles. Int} Pharm 214:9-12. 
210 Mader 
72. Lippacher A, Miiller RH and Mader K (2002) Semisolid SLN dispersions for topical 
application: Influence of formulation and production parameters on viscoelastic properties. 
Eur J Pharm Biopharm 53:155-160. 
73. Illing A and Unruh T (2004) Investigation on the flow behavior of dispersions of solid 
triglyceride nanoparticles. Int} Pharm 284:123-131. 
74. Souto EB, Wissing SA, Barbosa CM and Miiller RH (2004) Evaluation of the physical 
stability of SLN and NLC before and after incorporation into hydrogel formulations. 
Eur J Pharm Biopharm 58:83-90. 
75. Westesen K and Siekmann B (1997) Investigation of the gel formation of phospholipids 
stabilized solid lipid nanoparticles. Int J Pharm 151:35^5. 
76. Bunjes H, Westesen K and Koch MHJ (1996) Crystallization tendency and polymorphic 
transitions in triglyceride nanoparticles. Int J Pharm 129:159-173. 
77. Zimmermann E, Liedtke S, Miiller RH and Mader K (1999) H-NMR as a method 
to characterize colloidal carrier systems, Proc Inc Symp Control Rel Bioact Mater 26: 
591-592. 
78. Shahgaldian P, Da Silva E, Coleman AW, Rather Beth and Zaworotko MJ (2003) Paraacyl-
calix-arene based solid lipid nanoparticles (SLNs): A detailed study of preparation 
and stability parameters. Int J Pharm 253:23-38. 
79. Wissing, S A, Miiller RH, Manthei L and Mayer C (2004) Structural Characterization of 
QlO-Loaded Solid Lipid Nanoparticles by NMR Spectroscopy. Pharm Res 21:400^05. 
80. Ahlin P, Kristl J, Pecar S, Strancar J and Sentjurc M (2003) The effect of lipophihcity of 
spin-labeled compounds on their distribution in solid lipid nanoparticle dispersions 
studied by electron paramagnetic resonance. / Pharm Sci 92:58-66. 
81. S Liedtke, E Zimmermann, RH Miiller and K Mader (1999) Physical characterisation of 
solid lipid nanoparticles (SLN™), Proc Intl Symp Control Rel Bioact Mater 26:595-596. 
82. Kristl J, Volk B, Ahlin P, Gombac K and Sentjurc M (2003) Interactions of solid lipid 
nanoparticles with model membranes and leukocytes studied by EPR. Int } Pharm 
256:133-140. 
83. Miiller RH and Heinemann S (1994) Fat emulsions for parenteral nutrition IV: Lipofundin 
MCT/LCT regimens for total parenteral nutrition (TPN) with high electrolyte 
load. Int J Pharm 107:121-132. 
84. Klang SH, Parnas M and Benita S (1998) Emulsions as drug carriers — possibilities, 
limitations and future perspectives, in Emulsions and Nanosuspensions for the Formulation 
of Poorly Soluble Drugs, RH Miiller, S Benita and BBohm (eds.), Medpharm Scientific 
Publishers, Stuttgart, pp. 31-65. 
85. Davis SS, Washington C, West P and Ilium L (1987) Lipid emulsions as drug delivery 
systems. Ann N Y Acad Sci 507:75-88. 
86. Miiller RH and Runge SA (1998) Solid lipid nanoparticles (SLN) for controlled drug 
delivery, in Submicron Emulsions in Drug Targeting and Delivery, Benita S (ed.), Harwood 
Academic Publishers; Amsterdam, pp. 219-234. 
87. Mehnert W and Mader K (2001) Solid lipid nanoparticles: Production, characterization 
and applications. Adv Drug Del Rev 47:165-196. 
Solid Lipid Nanoparticles as Drug Carriers 211 
88. Westesen K (2000) Novel lipid-based colloidal dispersions as potential drug administration 
systems. Expectations and reality. Coll Polym Sci 278:608-618. 
89. Heydenreich AV, Westmeier R, Pedersen N, Poulsen HS and Kristensen HG (2003) 
Preparation and purification of cationic solid lipid nanospheres — Effects on particle 
size, physical stability, and cell toxicity. Int ] Pharm 254:83-87. 
90. Wissing SA, Kayser O and Miiller RH (2004) Solid lipid nanoparticles for parenteral 
drug delivery. Mv Drug Del Rev 56:1257-1272. 
91. Miiller RH, Radtke M and Wissing SA (2002) Nanostructured lipid matrices for 
improved microencapsulation of drugs. Int J Pharm 242:121-128. 
92. Bunjes H, Drechsler M, Koch MHJ and Westesen K (2001) Incorporation of the model 
drug ubidecarenone into solid lipid nanoparticles. Pharm Res 18:287-293. 
93. Sivaramakrishnan R, Nakamura C, Mehnert W, Korting HC, Kramer KD and Schafer- 
Korting M (2004) Glucocorticoid entrapment into lipid carriers — characterization by 
parelectric spectroscopy and influence on dermal uptake. / Control Rel 97:493-502. 
94. Lukowski G and Kasbohm (2001) Energy Dispersive X-ray Analysis of loaded solid 
lipid nanoparticles. J Proceedings — 28th International Symposium on Controlled 
Release of Bioactive Materials and 4th Consumer & Diversified Products Conference, 
San Diego, CA, United States, 516-517. 
95. Frederiksen HK, Kristensen HG and Pedersen M (2003) Solid lipid microparticle formulations 
of the pyrethroid gamma-cyhalothrin-incompatibility of the lipid and the 
pyrethroid and biological properties of the formulations. / Control Rel 86:243-252. 
96. Kristl J, Volk B, Gasperlin M, Sentjurc M and Jurkovic P (2003) Effect of colloidal carriers 
on ascorbyl palmitate stability. Eur J Pharm Sci 19:181-189. 
97. Olbrich C, Kayser O and Miiller RH (2002) Lipase degradation of Dynasan 114 and 116 
solid lipid nanoparticles (SLN) — effect of surfactants, storage time and crystallinity. 
Int J Pharm 237:119-128. 
98. Olbrich C, Kayser O and Miiller RH (2002) Enzymatic Degradation of Dynasan 114 
SLN — Effect of Surfactants and Particle Size. / Nanopar Res 4:121-129. 
99. Zimmermann E and Miiller RH (2001) Electrolyte- and pH-stabilities of aqueous solid 
lipid nanoparticle (SLN) dispersions in artificial gastrointestinal media. Eur J Pharm 
Biopharm 52:203-210. 
100. Mei Z, Li X, Wu Q, Hu S and Yang X (2005) The research on the anti-inflammatory 
activity and hepatotoxicity of triptolide-loaded solid lipid nanoparticle. Pharmacol Res 
51:345-351. 
101. Hu LD, Tang X and Cui FD (2004) Solid lipid nanoparticles (SLNs) to improve oral 
bioavailability of poorly soluble drugs. / Pharm Pharmacol 56:1527-1535. 
102. Bargoni A, Cavalli R, Zara GP, Fundaro A, Caputo O and Gasco MR (2001) Transmucosal 
transport of tobramycin incorporated in solid lipid nanoparticles (SLN) after duodenal 
administration to rats. Part II. Tissue distribution. Pharmacol Res 43:497-502. 
103. Cavalli R, Bargoni A, Podio V, Muntoni E, Zara GP and Gasco MR (2003) Duodenal 
administration of solid lipid nanoparticles loaded with different percentages of 
tobramycin. / Pharm Sci 92:1085-1094. 
212 Mader 
104. Zara GP, Bargoni A, Cavalli R, Fundaro A, Vighetto D and Gasco MR (2002) Pharmacokinetics 
and tissue distribution of idarubicin-loaded solid lipid nanoparticles after 
duodenal administration to rats. J Pharm Sci 91:1324-1333. 
105. Zara GP, Cavalli R, Bargoni A, Fundaro A, Vighetto D and Gasco MR (2002) Intravenous 
administration to rabbits of non-stealth and stealth doxorubicin-loaded solid 
lipid nanoparticles at increasing concentrations of stealth agent: Pharmacokinetics and 
distribution of doxorubicin in brain and other tissues. / Drug Targ 10:327-335. 
106. Chen D, Lu W, Yang T, Li J and Zhang Q (2002) Preparation and characterization of 
long-circulating solid lipid nanoparticles containing paclitaxel. Yixueban 34:57-60. 
107. Miiller RH, Radtke M and Wissing SA (2002) Solid lipid nanoparticles (SLN) and nanostructured 
lipid carriers (NLC) in cosmetic and dermatological preparations. Adv Drug 
Del Rev 54(Suppl 1):S131-S155. 
108. Wissing S A and Miiller RH (2002) The influence of the cry stallinity of lipid nanoparticles 
on their occlusive properties. Int} Pharm 242:377-379. 
109. Wissing SA and Miiller RH (2003) The influence of solid lipid nanoparticles on skin 
hydration and viscoelasticity — in vivo study. Eur J Pharm Biopharm 56:67-72. 
110. Wissing SA and Miiller RH (2001) Solid lipid nanoparticles (SLN) — a novel carrier for 
UV blockers. Pharmazie 56:783-786. 
111. Maia CS, Mehnert W, Schaller M, Korting HC, Gysler A, Haberland A and Schafer- 
Korting M (2002) Drug targeting by solid lipid nanoparticles for dermal use. / Drug 
Targ 10:489-495. 
112. Rudolph C, Schillinger U, Ortiz, A, Tabatt K, Plank C, Miiller RH and Rosenecker J 
(2004) Application of Novel Solid Lipid Nanoparticle (SLN)-Gene Vector Formulations 
Based on a Dimeric HIV-1 TAT-Peptide in vitro and in vivo. Pharm Res 21:1662-1669. 
113. Gasco MR, Zara GP, Saettone MF and PCT Int. Appl. (2004) Pharmaceutical compositions 
suitable for the treatment of ophthalmic diseases. Patent application WO 
2004039351. 
114. Videira MA, Botelho MF, Santos AC, Gouveia LF, Pedroso De Lima JJ and Almeida AJ 
(2002) Lymphatic uptake of pulmonary delivered radiolabeled solid lipid nanoparticles. 
/ Drug Targ 10:607-613. 
115. Stevens PJ, Sekido M and Lee RJ (2004) Synthesis and evaluation of a hematoporphyrin 
derivative in a folate receptor-targeted solid-lipid nanoparticle formulation. Anticancer 
Res 24:161-165. 
116. Peira E, Marzola P, Podio V, Aime S, Sbarbati A and Gasco MR (2003) In vitro and 
in vivo study of solid lipid nanoparticles loaded with superparamagnetic iron oxide. 
/ Drug Targ 11:19-24. 
117. Wang JX, Sun X and Zhang ZR (2002) Enhanced brain targeting by synthesis of 3',5'- 
dioctanoyl-5-fluoro-2'-deoxyuridine and incorporation into solid lipid nanoparticles. 
Eur } Pharm Biopharm 54:285-290. 
118. Zara GP, Cavalli R, Bargoni A, Fundaro A, Vighetto D and Gasco MR (2002) Intravenous 
administration to rabbits of non-stealth and stealth doxorubicin-loaded solid 
lipid nanoparticles at increasing concentrations of stealth agent: Pharmacokinetics and 
distribution of doxorubicin in brain and other tissues. / Drug Targ 10:327-335. 
10 
Lipidic Core Nanocapsules as New 
Drug Delivery Systems 
Patrick Saulnier and Jean-Pierre Benoit 
A new generation of controlled size Lipidic core NanoCapsules (LNC) is presented 
with respect to their simple formulation, interfacial characteristics, pharmacokinetic 
and biodistribution properties. We describe their ability to load and release 
hydrophobic drugs. 
1. Introduction 
The ultimate goal of therapeutics is to deliver any drug at the right time in a safe 
and reproducible manner to a specific target at the required level. A great deal 
of effort is currently made to develop novel drug delivery systems that are able 
to fulfil these specifications. Among them, nanoscale drug carriers appear to be 
promising candidates. Colloidal carriers are particularly useful because they can 
provide protection of a drug from degradation in biological fluids and promote 
its penetration into cells. However, because the body is so well equipped to reject 
any intruding object, for the materials to stand any chance of success within this 
hostile yet sensitive environment, they must be chosen very carefully. In particular, 
attention has to be turned to the composition of the surface of colloidal drug 
carriers.1 Indeed, their clearance rate from the circulatory system is determined by 
their uptake by the mononuclear phagocytic system (MPS), which in turn depends 
on their physico chemical surface characteristics. In order to enhance circulation 
time, steric protection of various nanoparticulate drug carriers can be achieved by 
the presence of hydrophilic and flexible polymers to their surface. In the search 
213 
214 Saulnier & Benoit 
for injectable, biocompatible and long-circulating systems, many colloidal systems 
have been evaluated. 
Different kinds of vectors can be used. For example, molecular vectors where 
the drug is complexed or associated to a transport molecule are currently used. 
Many vectors are also constituted by viruses or hybrid viruses, following the modification 
of their genomes in order to avoid the possibility of replication. In this 
way, they are used as gene delivery systems. However, we will focus on non viral 
vectors in this chapter. They are always formulated using soft physico chemical 
methods, by taking advantage of macromolecular self-assembly properties at the 
colloidal state in order to produce well-controlled particles. The number of required 
biological and physico chemical properties of these systems is high in order to formulate 
operant vectors. One of the most important specifications of these systems 
is the biocompatibility and biodegradability of each component that needs to be 
chosen carefully from a restricted list of molecules. Secondly, they need to be well 
constructed in terms of size and interf acial properties, in order to constitute stealthy 
systems that will not be phagocyted by the MPS and consequently will have the 
longest residence time in blood. 
We should not forget that such vectors exist biologically. Low density lipoproteins 
(LDL) are interesting systems possessing many of the required specifications. 
Unfortunately, their extraction, purification or reconstitution is still a challenge 
with strong physico chemical problems to solve. No convenient common solvent 
of proteins and lipids exists in order to reconstitute a similar supra-molecular framework. 
Consequently, we have to keep in mind a formulation of nanoparticles with 
biomimetic properties to those related to LDL as close as possible. 
We would now like to describe a novel class of nanoparticles (Lipidic core 
NanoCapsules:LNC) formulated without organic solvents with biocompatible and 
biodegradable molecules.2 We will see that after modification of the composition, 
we can control their size without difficulty in the 10-200 nm range, with a 
monomodal and narrow size distribution. 
Initially, we suggest describing the LNC formulation following some particular 
auto-organizational properties of Poly Ethylene Glycol (PEG)-like surfactants, 
induced by several emulsion-phase inversions in which they are incorporated. We 
will particularly emphasize the different physical methods that determine the characterization 
of the final structure of LNC, as well as their stability in suspensions. 
Then, we will describe strong correlations between their stealthy properties in blood 
and structural characteristics, mainly size and interfacial properties. In specific, we 
have evaluated the activation of the complement system in an original in vitro 
model. These nanocapsules are devoted to the encapsulation of drugs that need to 
be dispersed in their oily core. As a proof that the concept works, we will describe 
the ability of LNC to encapsulate and release simple lipophilic molecules, ibuprofene 
and amiodarone, in the last paragraph. 
Lipidic Core Nanocapsules as New Drug Delivery Systems 215 
2. Lipidic Nanocapsule Formulation and Structure 
2.1. Process 
The first step of the process consists of the formulation of a stable emulsion characterized 
by its oily phase (O), aqueous phase (W) and finally its surfactants 
mixture (S). 
Due to the complexity of the mixture, the brand names will be used throughout 
the following text. It is important to note that no organic solvent or mediumchain 
alcohols are used in the formulation. All these molecules are known to be 
biocompatible and biodegradable. This indicates that the lack of residual toxicity 
can guarantee the safe use of LNC for human administration. Solutol® is mainly 
comprised of 12-hydroxystearate of PEG 660 that corresponds to a hydrophilic surfactant 
(HLB = 11). The lecithin used is a mixture of hydrophobic phospholipids. 
The main compounds of each phase are reported in Table 1. 
The beginning of the formulation (see Fig. 1) corresponds to a magnetic stirring 
of all the components for which the proportions will be defined later, with a 
gradual rise in temperature from room temperature to 80°C at a rate of 4°C/min, 
leading to an W/O emulsion characterized by low conductivity. The system is 
Table 1 Compounds used in the LNC formulation. 
S • Solutol® HS-15:12-hydroxysterarate of PEG 660 and PEG 660 (low content) 
• Lipoid®: lecithin 
O • Labrafac®: triglycerides (C8-C10) 
W • Purified water 
• NaCl 
Fig. 1. Emulsion-phase inversion induced by temperature changes and the principle of 
LNC formulation. 
216 Saulnier & Benoft 
cooled from 80 to 55°C (4°C/min), leading to an O/W emulsion characterized by 
its high conductivity. Between these two kinds of emulsion, a transition zone called 
the Phase Inversion Zone (PIZ) is defined where the system is known to be in 
bicontinuous states.23 
In order to provide appropriate and optimal interfacial properties to the wateroil 
interfaces, the formulation typically requires three temperature cycles across the 
PIZ. The system is stopped at a temperature corresponding to the beginning of the 
PIZ, just before performing a final, fast-cooling dilution process in cold water (2°C). 
This second step of the formulation leads to LNC in suspension in an aqueous phase. 
The interfacial rheology method developed in several papers demonstrates 
that the interfacial association of all the implicated molecules of the process is different 
from other commoner systems.4 Cohesion energy at the interface, as well 
as the interaction of the interfacial molecules with the adjacent phases, reaches a 
minimum for the concentrations used. We think that this particularity can explain 
why the system can be broken down in an ideal way during final dilution. The 
surfactants involved in the stabilization of the bicontinuous systems can easily 
leave the microemulsion in order to constitute the colloidal structures (LNC). 
It might be noted that temperatures corresponding to the PIZ are much too high 
to decline this method to the simple encapsulation of thermo-sensitive molecules. 
Fortunately, we have shown that the electrolyte concentration (NaCl) strongly influences 
the location of PIZ on the temperature scale. When we increase the electrolyte 
concentration, we decrease the PIZ temperature to reach acceptable levels. 
2.2. Influence of the medium composition 
Obviously, the presence or not of LNC strongly depends on the composition of 
the system reported in Fig. 2(a) as a pseudo-ternary diagram.5 Each point corresponds 
to strictly similar formulation processes and the entire diagram describes 
the appropriate feasibility zone. 
It should be noticed that the optimal formulation corresponds to w /w concentration 
of around 20% for the oil phase, 60% for the water phase and 20% for 
Solutol®. In the zone corresponding to the LNC formulation, a statistical model 
is applied in order to approximate the influence of the composition on the size 
distribution measured by the dynamic light scattering method. 
Polynomial interpolations between well-controlled points are performed. The 
corresponding results are reported in Fig. 2(b) where different iso-size curves are 
presented. The same procedure was applied to the size variation coefficients. These 
two curve beams are powerful tools, allowing an optimized formulation to be 
found, once a given and reproducible size distribution is elaborated just by tuning 
the composition. 
Lipidic Core Nanocapsules as New Drug Delivery Systems 21 7 
(a) (b) 
Fig. 2. Feasibility diagram of LNC. a: zone of favorable formulation; b: iso-size curves in 
the favorable zone. 
Fig. 3. Schematic representation of LNC. 
It is important to note that LNC have demonstrated very good freeze-drying 
and stability characteristics in storage conditions for several months, as determined 
by DSC measurements,6 confirming the structure presented in Fig. 3. 
LNCs are constituted of a lipidic core surrounded by a surfactant shell, where 
lecithin is located in the inner part of the shell and the Solutol® in the outer part. 
2.3. Structure and purification of the LNC by dialysis 
Considering that in the biological environment of the blood stream, the particles 
interact strongly with various interfaces, one possible model for studying the interfacial 
behavior of these particles is their spreading at the air-water interface. Classically, 
the Langmuir balance was used to describe interfaces composed by simple 
21 8 Saulnier & Beno?t 
mixtures. The basic technique was the measurement of the surface pressure (7r)-area 
(A) isotherm, by determining the decrease in surface tension as a function of the 
area available for each molecule on the aqueous sub phase. This included the study 
of the monolayer formation, the compressibility of the interface, the mutual interactions 
of molecules in the monolayer, but also interactions with the sub-phase 
molecules across interfacial rheological measurements.7 Following this, these suspension 
spreading results were compared with zeta potential measurements. These 
studies8,9 clearly indicate that the mother suspension, just after dilution in cold 
water, is composed of 
• Stable nanocapsules as described before; these objects diffuse strongly in the 
aqueous phase after spreading at the air/water interface. 
• Unstable nanocapsules with similar size, but with a lower amount of phospholipids 
(Lipoid®) in the inner part of their shell. These capsules are not sufficiently 
robust to support the interfacial energies during spreading. Consequently, the 
components or fragments of the initial particles can be detected at the air-water 
interface. 
• Free PEG (minor component of the Solutol®) released from the outer part of 
the shell. 
It is obvious that the excess of PEG, as well as an important fraction of the 
unstable particles could be limited by dialysis. We will see in the next chapter an 
original investigation of these dialysis effects. 
2.4. Imagery techniques 
AFM images [Fig. 4(a)] were obtained after spreading the initial suspension of 
50 nm (±10 nm) LNC on a fresh mica plate, and then allowing a complete evaporation 
of the water at room temperature. A contact mode was applied with a contact 
force of 10 nN, as well as a non contact mode without modification of the related 
images. The particle shape looked like a cylinder, 2nm high and 275 nm wide, 
corresponding to a total volume similar to a 60 nm sphere. We demonstrate the 
deformation of LNC after water evaporation, but without fusion of the particles, 
something that often occurs with liposomes. 
Classical TEM images were taken of the covered copper grids, following staining 
with a 2% phosphotungstic acid aqueous solution. It is noted on Fig. 4(b) that 
the lateral diameters are relatively polydispersed in a 20-70 nm range. 
Fig. 4(c) corresponds to a cryo-TEM image (kindly provided by Olivier Lambert, 
IECB-UBS UMR CNRS 5471) where individualized LNC are detectable. It is 
important to note that this image was performed after a dialysis, followed by an 
appropriate dilution of the mother suspension. 
Lipidic Core Nanocapsules as New Drug Delivery Systems 219 
(b) TEM (c) Cryo-TEM 
<>,J 'HUE ^ > 
* "s ' * * 
*o * If 
Fig. 4. Visualization of LNC by (a) AFM, (b) TEM and (c) cryo-TEM. 
3. Electrical and Biological Properties 
3.1. Electro kinetic comportment 
The stable Lipid NanoCapsules (LNC) contain pegylated 12-hydroxy stearate, as 
well as free PEG in the outer part of the shell, which can be an important biological 
specification that we will describe latter. The distribution of PEG chains at the 
surface was determined by their electrokinetic properties. Thus, electrophoretic 
mobility was measured as a function of ionic strength and pH, for particles differing 
in sizes, dialysis effects, and the presence or not of lecithin in their shell. The study 
enabled us to find the isoelectric point (IEP) as well as the charge density (ZN) in 
relation to the dipolar distribution in the polyelectrolyte accessible layer (thickness 
1 A), by using soft particle electrophoresis analysis10 (see Fig. 5). 
This study showed that LNC presented electrophoretic properties conferred by 
PEG groups at the surface constituting dipoles that are able to interact with counter 
ions (H+, Na+) or water dipoles. 
The levels of IEP, ZN and 1/1 changed after dialysis, due to the removal of 
molecules that were poorly linked (mainly free PEG) at the outer part of the surface, 
allowing accessibility to the inner adjacent part of the shell. 
Water shell 
Fig. 5. Accessible layer to counter ions characterized by its thickness (1 A ) and its dipolar 
charge density (ZN). 
(a) AFM 
* 1 
m ! 
it i 
• '• 
:&&&* • • ' . , . ) A • •••: . : - ? .' 
•••-ViCS. . ' . i f , . 
I - " • • V o ^ -7 L—» J*^ •, ft". - •' 
n •:•%*. '.••-.•:?»••&* 
H -':' •''"•"' v ' • •' 
• .•.:.->.;.:- •• .. 
m • « " • • > • • •, 
.) urn 
220 Saulnier & Beno?t 
100 nm LNC presented the best-organized and the accessible part of the shell, 
compared with other sizes of LNC, before and after dialysis. Lecithin was found to 
be present in the inner part of the polyelectrolyte layer and was found to play a role 
in the disorganization of the outer part. Dialyzing LNC formulated with lecithin 
led to stable and well structured nanocapsules, ready for an in vivo use as a drug 
delivery system.11 
3.2. Evaluation of complement system activation 
Generally, after intravenous administration, nanoparticles (NP) are rapidly 
removed from the blood stream because they are recognized by cells of the MPS such 
as Kiipffer cells in the liver, or spleen and bone-marrow macrophages. However, a 
brush of PEG chains grafted on the surface is known to decrease the recognition of 
nanoparticles by the immune system after intravenous administration.12 One has 
demonstrated that a strong correlation prevails between the complement activation 
and the stealthy properties of LNC. 
Therefore, these properties were evaluated by measuring the degree of complement 
activation11 [CH50 technique and crossed Immunoelectrophoresis (C3 cleavage)] 
and the level of macrophage uptake, in relation to the organization of PEG 
chains, according to the electrokinetic properties of the LNC surface. These experiments 
were performed on 20, 50 and 100 nm LNC before and after dialysis. The 
CH50 technique is presented in Fig. 6. 
Nanoparticles are dispersed in human serum with sensitized erythrocytes. 
After incubation, lysis is evaluated by a classical spectrophotometric method. The 
measured absorbance is related to the consumption of complement proteins by 
particles. 
The main conclusions are that whatever the in vitro test, all LNC were not recognized 
by the non specific components of the immune system. It was probably due 
to the strong density of PEG chains at their surface. Furthermore, dialysis maintains 
a sufficiently high density of PEG and had no incidence on the complement 
consumption. 
4. Pharmacokinetic Studies and Biodistribution 
At first, the biodistribution of radiolabeled nanocapsules was studied by scintigraphy 
and y counting, after intravenous administration in rat whereby the 99mTc-oxine 
was incorporated in the lipid core and 125I labelled the shell of the nanocapsules.13 
Dynamic scintigraphic acquisition was carried out 3 hrs after administration and y 
activity in blood and tissues was followed for more than 24 hrs (see Fig. 7). 
An early half-disappearance time of about 47 ± 6 min was found for 125I and 
41 ± 11 min for 99mTc. These ranges of residence times were interesting for specific 
Lipidic Core Nanocapsules as New Drug Delivery Systems 221 
«—car"" 
Lysis of 
erythrocytes 
M, 
**CSf—. 
^B Sheep erythrocyte 
• Complement proteins 
M Amibody anti-sheep eryihrocyie No lysis of 
ervlhrocvtes 
Fig. 6. CH50 method for the evaluation of complement system activation. 
200 300 
Time (min) 
500 600 
Fig. 7. Evolution of radioactivity blood repartition after the intravenous administration of 
LNC expressed as a percentage of the injected dose. 
222 Saulnier & Benoit 
site delivery. Meanwhile, it appears that the length of the PEG chain (in this case, 
15 ethylene oxyde groups per molecule) should be increased to extend the vascular 
residence time. 
Recently, it has been shown that adding different DSPE-PEG to the system 
enhances the t1/2 values to several hours, depending on the concentration and the 
PEG length.14 t1/2 (half-life), MRT (Mean Residence Time) in blood and AUC (Area 
Under Curve) were evaluated by using [3H]-cholesteryl hexadecyl ether mixed with 
the lipid and the surfactant at the beginning of the formulation. 
The main conclusion was that the LNC formulated in this study compared 
advantageously with other nanoparticulate systems, particularly for their residence 
time in blood. Nanocapsule uptake by the different organs of rat was evaluated 
24 hrs after intravenous administration. It was shown that LNC deposited mainly 
in the liver and the spleen, but also in the heart, and the results were comparable 
to a liposome reference. 
5. Drug Encapsulation and Release 
5.1. Ibuprofene 
LNC were characterized for their suitability as an ibuprofene delivery device for 
pain treatments.15 After in vitro investigations, ibuprofene- loaded LNC were evaluated 
after intravenous and oral administration in rats. For each system, the carrier 
was evaluated through its potential antinociceptive efficiency. 
We present in Fig. 8, the release of ibuprofene in a phosphate buffer after 
its incorporation in LNC during formulation. For each case, LNC provide high 
ibuprofene loadings (95%). The main feature is an initial burst followed by a 
100 
CP" 80 
o^ 
W atex beads 
l : < 
r> % V 
- J1 
- • • • © • • • • • V - : 
• • . • • ; ' . . • • , , . « 
: WOK 
Fig. 8. Sizing studies of submicron bubbles. 
distribution is less than one micron and the mean size is less than 1 micron in 
diameter.14'15 
The images below depict a photomicrograph of MRX-815 bubbles alongside a 
photomicrograph of one-micron size latex microbeads. The bubbles are one micron 
in diameter and smaller. We found in our lab that the smallest bubbles are not 
well shown on the light microscopy due to limitations of the imaging technique. 
The sizing profile shows that there are bubbles up to approximately two microns 
in diameter, but more than 70% of the bubbles are smaller than one micron in 
size. 
16,17 
Investigators have demonstrated that ultrasound can be used to generate cavitation 
in an aqueous medium.18 Cavitation research has led to studies involving 
ultrasound-mediated clot lysis at a variety of frequencies.19"23 Furthermore, 
microbubbles and submicron-sized bubbles provide a nucleus at which cavitation 
can occur, thereby lowering the ultrasound energy requirements.24 
While intravenous administration with local application of ultrasound appears 
to be effective for sonothrombolysis in both pre-clinical and clinical models, applications 
using an infusion catheter are also being investigated.25 It is believed 
that submicron-sized bubbles and ultrasound-mediated cavitation are able to 
affect the thrombus architecture by increasing permeability through the thrombus 
matrix, thereby improving accessibility and the penetration of thrombolytic 
enzymes to more efficiently lyse clots. Studies by Francis et a/.,26'27 demonstrated 
that ultrasound alone increased the spacing between fibrin strands in 
clots, presumably improving the penetration of lytic enzymes, such as t-PA, into 
the clot. 
By way of explanation, when bubbles are insonified, these bubbles can oscillate 
in response to the acoustic pressure wave. If driven with a sufficient acoustic 
pressure, the rapid expansion and contraction of the bubble will result in local 
234 Ungeretal. 
velocities at the bubble surface on the order of hundreds of meters per second. 
If the expansion of the bubble is large enough, the bubble will become unstable, 
resulting in the destruction of the bubble into smaller fragments.28 The rapid 
oscillation of the bubble in response to an acoustic pulse is referred to as "cavitation". 
Bubbles undergoing this violent expansion and contraction produce liquid 
jets, local shock waves, and free radicals. Although the exact mechanism is still 
being studied, the effect of cavitating bubbles has been demonstrated to have several 
effects on the surrounding tissues, including the poration of cell membranes 
resulting in enhanced membrane permeability (sonoporation) or the disruption 
of local thrombus. Thus, the combination of ultrasound with microbubbles has 
potential applications in blood clot dispersion and local drug delivery to treat 
cardiovascular disease, cancer, and diseases of the central nervous system. The 
figure below shows individual images from ultra-high speed videomicroscopy of 
a single bubble. The bubble is shown in the resting state on the far left hand side 
of the figure. The bubble expands after the application of the ultrasound pulse, 
then collapses and fragments. The daughter bubbles expand and collapse again, 
leaving behind small nano-sized fragments.29 Localized activation of bubbles with 
ultrasound can be used for a number of different medical applications including 
SonoLysis. 
Whereas in diagnostic ultrasound contrast imaging where there is an r6 dependence 
between size and ultrasound reflection for therapy, it is advantageous to 
have much smaller bubbles. As shown in Fig. 10, when bubbles are cavitated by 
ultrasound, they may undergo a relatively greater increase in the expansion ratio 
ri/ror where r^ is the maximum size for the radius of the bubble after insonation, and 
r0 — the initial resting radius.31 The relative expansion with insonation is greatest 
for the smallest diameter submicron-sized bubbles. This conceivably results in a 
more effective cavitational force, and hence more efficient lysis of thrombi. 
Another effect of ultrasound on microbubbles which has the potential to be utilized 
therapeutically is the use of acoustic radiation force to selectively concentrate 
microbubbles at a target site.32'33,34 Microbubbles driven with ultrasound, experience 
radiation force in the direction of ultrasound wave propagation.35 Pulses of 
t % '2 IV \ ,' V y. 
Fig. 9. In the images above, a single 3 (im bubble is shown (far left) in the resting state. 
Insonation with a single pulse of ultrasound energy causes the bubble to expand, collapse, 
and fragment, yielding nanometer-sized fragments. As the bubbles expand and collapse, they 
generate a local Shockwave that can be used therapeutically. Reproduced with permission 
from Chomas et (A., Appl Phys Lett, 2000.30 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 235 
Fig. 10. The relationship between nanobubbles' size at resting state and expansion ratio 
under insonation. Reproduced with permission of D. Patel et a\., IEEE Ultrasonics, Ferroelectrics, 
and Frequency Control. In press. 
many cycles can deflect resonant microbubbles over distances in the order of millimeters. 
Thus, it may be possible to bring microbubbles circulating in the blood 
pool into contact with targeting sites on a blood vessel wall, in a region selected 
by the positioning of the ultrasound beam. This effect has been demonstrated to 
increase the retention of microbubbles at a target site over an order of magnitude.3 6 
In addition to favorable acoustic characteristics, submicron-sized bubbles have 
other potential advantages for therapy, compared with larger-sized microbubbles. 
The smaller bubbles may penetrate a clot more easily and may have better biodistribution 
characteristics for targeting. 
The pictorial representation below (Fig. 11) is the hypothetical mechanism of 
action for MRX-815 bubbles flowing through the vasculature in association with 
Fig. 11. It is hypothesized that when submicron-sized bubbles are injected systemically, 
some will aggregate on the thrombus, and due to their small size, work into the clot. When 
the bubbles cavitate, the kinetic energy disperses the clot, both from its periphery, and due 
to the fact that bubbles are able to penetrate the clot from within. 
236 Ungeretal. 
a thrombus. Ultrasound could cause cavitation of the bubbles, transferring their 
dispersive energy to the clot and dispersing the clot safely and painlessly. Particle 
sizing studies of the effluent from in vitro studies of SMB-assisted sonothrombolysis 
have shown that the particles are submicron in size.37 
The figures below show the experimental set-up used in our lab for a flow 
through phantom for testing sonothrombolysis, and then treatment of a clot in 
the phantom. The clot was exposed to 1 MHz ultrasound and tissue plasminogen 
activator (t-PA), followed by an infusion of MRX-815 microbubbles. As shown in 
the figures, after 40 min of treatment there is near complete resolution of the clot. 
The graph below shows the results from a series of clots exposed to t-PA, 
t-PA + ultrasound and t-PA + ultrasound + MRX-815 bubbles in our lab. Note 
that the greatest reduction of thrombi was in the group exposed to bubbles. 
, - 
C 
Fig. 12. Above a schematic of the experimental set-up: (A) the clot pre-treatment, (B) after 
32 min of treatment, (C) after 40 min of treatment. The clot was 96% dissolved. 
80.00 
70.00 
60.00 
••2 50.00 
Z 40.00 
2. 
« 30.00 
s? 
20.00 
10,00 
0.00 
Saline US t-PA t-PA, SMB, SMB, 
US US US. t-PA 
Fig. 13. SMB = Bubbles. 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 237 
4. Clinical Studies 
Vascular thrombosis is a major cause of death in industrialized countries, responsible 
for myocardial infarction, stroke and peripheral arterial occlusions.38 In 
addition, deep vein thrombosis (DVT), which afflicts one in twenty Americans during 
their lifetime,39'40 may also be an application for sonothrombolysis. 
ImaRx completed a Phase I/II clinical trial in thrombosed dialysis grafts for 
the purpose of preliminary feasibility and safety for sonothrombolysis treatment 
of clotted grafts. Initial studies in thrombosed dialysis grafts provided a venue to 
evaluate the principle of sonothrombolysis in vascular thrombosis. As such, clinical 
trial efforts will move forward to address the treatment of stroke, peripheral arterial 
occlusions (PAO) and deep vein thrombosis (DVT). 
Below are examples shown from clinical trials for sonothrombolysis in dialysis 
grafts and DVT. The examples are not an indication that all sonothrombolysis 
treatments will have similar outcomes. 
Images from a venogram in a patient with DVT showed that the patient was 
administered bubbles via infusion catheter into the popliteal vein over a period of 
1 hr, while ultrasound was applied across the skin. No thrombolytic drug such as 
t-PA was administered. Clinically, this particular patient had marked reduction in 
pain post-treatment with sonothrombolysis. 
Stroke is the third most common cause of death, after heart disease and cancer 
in North America. It incurs far more expenses than any other diseases due to its 
long term disability.41 In the US, stroke accounts for over $50 billion each year 
to the health care system.42 The only approved pharmacologic therapy to help 
restore blood flow in stroke patients is t-PA (Activase®). Less than 5% of patients 
are treated with t-PA due to concerns over bleeding and the risk relative to the 
benefit.43 Encouraging results have been obtained, however, in human studies with 
ultrasound and t-PA, and most recently, with ultrasound + t-PA + microbubbles. 
Fig. 14. The j:iyio.i a:n on the left is of a clotted dialysis graft. Very little contrast enters the 
graft as it is filled with clot. The image on the right, post-bubble treatment, shows complete 
opacification of the graft due to successful dissolution of thrombosis by sonothrombolysis. 
238 Ungeretal. 
Fig. 15. On the pre-treatment image (left), there is complete occlusion of the superficial 
femoral vein (SFV). Collateral veins are seen carrying the blood flow that would normally 
be carried by the SFV. Post-treatment, there is good flow in the SFV and much less flow is 
seen in the collateral vessels due to the increased flow in the SFV. 
Dr. Andrei Alexandrov from the University of Texas in Houston led a study of 
ultrasound + t-PA in acute ischemic stroke.44,45 In this study, 126 patients were randomized 
prospectively to receive either a 1 hr infusion of t-PA at a dose of 0.9 mg/kg 
alone, or t-PA plus 2 hrs of continuous trans-cranial Doppler (TCD) ultrasound 
applied through the temporal window where the skull is thinnest and most easily 
penetrated by ultrasound. Of the 63 patients treated with t-PA alone, there was a 
13% recanalization rate of the intra-cranial circulation at 2 hrs.46,47 In the same number 
of patients receiving t-PA + ultrasound, there was a highly significant increase 
in recanalization to 38% at two hours, indicating that ultrasound-mediated therapy 
aided in thrombus dispersion. 
Dr. Carlos Molina, from Barcelona, Spain, conducted a similar study but with 
microbubbles.48 The addition of microbubbles enhances the cavitational nuclei 
with a decrease in power requirements. Dr. Molina's study demonstrated that 
the recanalization rate increased impressively to 55%.49 In this study, Dr. Molina 
administered three doses of Levovist®, a microbubble agent comprised of air-filled 
galactose microparticles. Dr. Molina's pioneering work has demonstrated the utility 
of using bubbles in conjunction with ultrasound to improve the clinical outcome 
of acute stroke. ImaRx is currently moving MRX-815 into stroke treatment 
trials. 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 239 
SMB-assisted sonothrombolysis therapy could move beyond the current clinical 
regimens by eliminating the thrombolytic agent. Pre-clinical trials in both canine 
and porcine models have been encouraging.50'51-52 Human studies will be conducted 
to determine if lipid-coated bubbles will improve recanalization rates in patients 
treated with this new ultrasound-mediated paradigm. 
5. Blood Brain Barrier 
Poor transport into the CNS is an obstacle to effectively treat diseases including 
brain tumors, Alzheimer's and other neuro-degenerative diseases. There are two 
principal barriers to drug transport into the CNS: (a) the blood brain barrier (BBB) 
and (b) the ABC transporters, ABCC1 and ABCB1. 
Unlike the rest of the body, the capillary foot processes of the cerebral endothelial 
cells are tight, preventing peptides and macromolecules from leaking through 
to the brain.53 Although the BBB may be permeable to selected ions and small 
molecules, ABCB1, also known as the P-glycoprotein, acts to remove the molecules 
by a drug-efflux system before they enter the brain. Several different strategies have 
been developed to overcome these limitations.54 One approach to drug delivery to 
the brain is by the transient opening of the BBB. 
Hypertonic solutions containing mannitol, which act by shrinking the endothelial 
cells when co-administered with drugs, have been shown to result in enhanced 
cerebral drug uptake.55,56 However, to cause minimum side effects, it is essential for 
the therapy to be regional and localized. Recently, Hynynen et al.57 have shown that 
the BBB can be transiently opened using ultrasound and microbubbles (Illustrated 
in Fig. 16). When bubbles were administered intravenously and focused ultrasound 
was applied across the intact skull, the BBB could be reversibly opened, permitting 
passage of hydrophilic low molecular weight molecules such as gadolinium-DTPA, 
and macromolecules such as fluorescently labeled albumin (Fig. 17) into the CNS.58 
The permeability resolved over a period of hours without damage to the neurons. 
Similar studies have been performed in a porcine model showing that nonfocused 
ultrasound with microbubbles can be used to open the BBB.59 Figure 18 
shows increased dye deposition in the cerebral tissue. 
Introduction of microbubbles as the cavitation nucleus prior to the application 
of ultrasound, lowered the energy needed to open the BBB, thereby lowering 
the bioeffects of ultrasound.60 Using this technique, large biomolecules such as 
horseradish peroxidase (a 40 kDa protein) have been shown to pass through the 
BBB with minimal damage to the brain tissues.61 
It can be envisaged that drugs (small or macromolecules) bound to the 
microbubbles would function as a more efficient drug delivery vehicle, since these 
240 Unger et al. 
A 
/ , 
0 <(S 
/ / 
/ MlLI'JCUDblO 
Nii'iadrofltil 
B 
UltMSOUll'l 
*@ 
c 
' < * 
# • « 
Fig. 16. Cartoon representation of hypothesized ultrasound mediated drug delivery to the 
brain. (A) Cerebral capillaries with tight endothelial junctions prevent passage of molecules 
(including microbubbles and nanoparticles) into the brain. (B) Ultrasound is applied to 
the skull through the temporal window where the skull is thinnest (inset), cavitating the 
microbubbles and opening up the endothelial junctions. (C) Therapeutic agents may now 
pass through the opened junctions. 
Location 
1 
2 
3 
4 
Pressure 
amplitude 
values 
4.7 MPa 
2.3 MPa 
3.3 MPa 
1.0 MPa 
Fig. 17. Tl-weighted MR images of rabbit brain after treatment shows contrast enhancement 
at 4 locations (arrows), coronal image across focal plane. Reproduced with permission 
from Hynynen et ah, Radiology. 
would provide the cavitation nuclei and the drug payload in one entity, circumventing 
the co-administration of drug and microbubble. In such instances, the drug 
could be (a) bound to the lipid membrane (hydrophobic drugs), (b) bound to the 
charged lipids on the surface (gene delivery), or (c) buried in the interior in an oily 
layer of a droplet (hydrophobic drugs) (Fig. 19). Furthermore, (d) these drug loaded 
bubbles or droplets may have the potential to be targeted to a specific site in the 
brain by surface ligands. 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 241 
ug/g tissue 
30 
25 
20 
15 
10 
P=0.83 
30 
25 
20 
15 
10 
P=0.006 
Untreated Ultrasound Untreated Ultrasound + MB 
Fig. 18. Control pigs and pigs treated with ultrasound alone showed no difference in Evan's 
blue uptake. There was a significant difference in uptake when microbubbles were used in 
conjunction with ultrasound. Adapted from Porter et ah,} Am Soc Echocardiogr. 
Fig. 19. Different ways that bubbles or droplets may be able to transport drugs. Drugs may 
be (a) bound or embedded in the lipid membrane, (b) bound to the surface charges of the 
phospholipid membrance (c) buried in the oil in a droplet (d) targeting ligands can be incorporated 
onto the membrance. 
This technology of activation with ultrasound and microbubbles has the potential 
to also be used in the drug discovery process. By exposing cultured neurons 
to drugs, ultrasound and bubbles, high concentrations of the drug may be able to 
deliver to the cells without damaging them. This can potentially be used to screen 
neurons for new therapeutic compounds. 
242 Unger et al. 
Potential CNS diseases amenable to treatment with 
submicron bubble delivery and classes of drugs 
Disease Drugs 
Alzheimer's Disease and 
other neurodegenerative 
disease, seizures and 
psychiatric disorders 
Primary and Secondary 
(metastases) Brain Tumors 
Stroke, brain ischemia 
Infection, e.g., AIDS 
Low molecular weight therapeutics with poor 
delivery to CNS, proteins, gene-based therapeutics. 
Low molecular weight therapeutics with poor 
delivery to CNS, proteins, genetic drugs. Radiation 
sensitizers. 
Cavitation nuclei to augment sonothrombolysis, 
either with or without use of thrombolytic 
agent. Delivery of oxygen with microbubbles. 
Improvement of cerebral perfusion with 
microbubble-enhanced sonication. Delivery of 
anti-oxidants and growth factors. 
Delivery of anti-infectives, anti-retrovirals to 
CNS. 
6. Drug Delivery 
In the foregoing sections, we discussed activating the bubbles or using them 
in conjunction with ultrasound-mediated processes (e.g. microbubble mediated 
sonothrombolysis to enhance the local activity of the drug such as t-PA), or that 
the availability of a drug may be increased, e.g. by opening the blood brain barrier. 
In this section, we will discuss evaluating drug-carrying microbubbles for drug 
delivery. 
6.1. Targeted bubbles 
As preliminary studies to demonstrate feasibility of using targeted bubbles as 
potential drug delivery agents, two different targeted bubbles were prepared 
using a mixture of DPPC, DPPE-PEG5000 and DPPA, as well as different oils 
and perfluorocarbons using a mixture of DDFP and n-perfluorohexane. In one 
study, a bioconjugate ligand targeted to the am, An integrin was synthesized by 
solid phase peptide methodology.62 Briefly, the bioconjugate, lipids, biocompatible 
drug, perfluoropropane were combined into a mixture and bubbles prepared 
by shaking the vials at approximately 4200 rpm. The size of the targeted bubbles 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 243 
Fig. 20. Intravital microscopy demonstrating adherence of targeted microbubble to thrombus. 
Picture on the right is a graphic representation outlining the location of bound microbubbles 
on thrombus. Reproduced with permission from Schumann et ah, Investi Radiol. 
was approximately 2 fim, as measured by light obscuration measurements on a 
Particle Sizing Systems Model 470 sizer (Particle Sizing Systems, Santa Barbara, 
Calif.). Bubbles were injected into a mouse model where thrombi were previously 
formed in the cremasteric arterioles and venules. Fluorescent imaging revealed 
binding of the targeted bubbles to the thrombi in both arterioles and venules. 
Figure 20 demonstrates the utility of a targeted bubble. 
Similarly, targeted bubbles were used in a HUVEC cell culture model. Briefly, 
bubbles with a targeting ligand directed to a^ft receptors on HUVEC cells were 
(a) 
O 
 1 /xm), which allows them to 
be administered intravenously without any risk of embolization. According to the 
process and the composition used in the preparation of nanoparticles, nanospheres 
255 
256 Gref & Couvreur 
Fig. 1. (A) Schematic representation of the nanocapsule structure; (B) Morphological 
appearance of a nanocapsule with an oily core (transmission electron microscopy after freeze 
fracture). 
or nanocapsules can be obtained. Nanospheres are matrix systems in which the drug 
is dispersed within the polymer throughout the particle. Contrarily, nanocapsules 
are vesicular or "reservoir" (heterogenous) systems, in which the drug is essentially 
confined to a cavity surrounded by a tiny polymeric membrane (Fig. 1). As in the 
case of nanospheres, depending on their physicochemical properties and composition, 
the drug may adsorb onto the surface as well as being included in the central 
core of nanocapsules. Therefore, drug localization is an important parameter in the 
characterization of nanocapsule preparations. 
The nanocapsule core may be acqueous or composed of a lipophilic solvent, 
usually an oil. In order to achieve good drug loading, the core materials are chosen 
among the good solvents for the drug.1 Expected advantages of confining the drug 
within a central cavity are: (a) burst effect may be avoided; (b) the drug is not 
in direct contact with tissues and therefore irritation at the site of administration 
could be reduced, and (c) the drug may be better protected from degradation both 
during storage and after administration. One of the advantages of nanocapsules 
over nanospheres is their low polymer content and a high loading capacity for 
lipophilic drugs. 
Nanocapsules can either be obtained by interfacial polymerization of 
monomers or from preformed polymers. In the former, the molar mass of the coating 
polymer will depend on the preparation conditions and even on the drug used, 
whereas in the latter, it is determined at the outset. Polymerization of monomers 
may lead to a covalent linkage between the polymer and the drug. To date, all the 
methodologies described for preparing nanocapsules involve the preparation of 
emulsions. Oil-in-water (O/W) emulsions lead to the formation of nanocapsules 
with an oily core, suspended in water. Water-in-oil (W/O) emulsions lead to the 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 257 
obtention of nanocapsules with an acqueous core, suspended in oil. More recently, 
nanocapsules with an acqueous core suspended in an acqueous medium were also 
obtained. 
Nanocapsule technology and their pharmaceutical applications will be further 
discussed according to the method of obtaining the polymeric wall (polymerization 
in situ or preformed polymer) and whether the core is acqueous or oily. 
2. Preparation 
2.1. Nanocapsules obtained by interfacial polymerization 
The advantage of obtaining nanocapsules by interfacial polymerization is that the 
polymer is formed in situ, allowing the polymer membrane to follow the contours of 
the inner phase of an O/W or W/O emulsion, thus entrapping drugs with high loadings. 
However, because reactive monomers are used, unwanted chemical reactions 
may occur between the drug and the monomer, before or during the polymerization 
process. 
The preparation of nanocapsules by polymerization requires a fast polymerization 
of the monomers at the interface between the organic and the acqueous phase of 
the emulsions. Alkylcyanoacrylates, which polymerize within seconds, have been 
proposed for the preparation of both oil- and water-containing nanocapsules. Their 
polymerization is initiated by hydroxyl ions either from the equilibrium dissociation 
of water or by nucleophilic groups of any compound in the polymerization 
medium.2 
2.1.1. Oil-containing nanocapsules 
The oil-containing nanocapsules are suitable for the encapsulation of the lipophilic 
and oil-soluble compounds. They are generally obtained by interfacial polymerization 
of alkylcyanoacrylates, after preparing a very fine oil-in-water emulsion 
with an additional water-miscible organic solvent such as ethanol or acetone.3'4 
These solvents serve as vehicles for the monomers, and also help to disperse the 
oil as very small droplets in the acqueous phase, which contains a hydrophilic 
surfactant. Indeed, as pointed out by Gallardo et al.,5 the organic solvents must 
be completely water-miscible, so that the formation of small enough oil droplets 
occurs spontaneously, while the solvent is diffusing towards the acqueous phase 
and the water is diffusing toward the organic phase. Meanwhile, the polymerization 
of the monomer induced by the contact with hydroxyl ions from the water phase 
must be swift to allow efficient formation of the polymer envelope around the oil 
droplet, thus achieving effective encapsulation of drugs. Generally, particles with 
258 Gref & Couvreur 
sizes ranging between 250 and 300 nm, depending on the experimental conditions, 
were obtained.5,6 
In a general procedure of nanocapsule preparation, the oil, the monomer, and 
the biologically active compound are dissolved together in the water-miscible 
organic solvent to prepare the organic phase.3-9 This organic phase is then injected 
via a cannula, under strong stirring, into the acqueous phase containing water and 
a hydrophilic surfactant. The nanocapsules are formed to give a milky suspension 
immediately. The organic phase is then removed under reduced pressure and the 
nanocapsules are purified by ultracentrifugation. Depending on the density of the 
oil forming the core, nanocapsules will concentrate either as a pellet at the bottom 
of the ultracentrifuge tubes or as a floating layer at the top of the tubes. 
A wide range of oils is suitable for the preparation of nanocapsules, including 
vegetable or mineral oils and pure compounds such as ethyl oleate and benzyl 
benzoate. The criteria for selection are the absence of toxicity, lack of affinity for 
the coating polymer, the absence of risk of degradation of the polymer, and a high 
capacity to dissolve the drug that is entrapped. Generally, Miglyol® is used to 
form the core of the nanocapsules.3-7,9,10 Lipiodol® and benzyl benzoate have also 
been successfully used to form nanocapsules.4 Soluble surfactants were chosen 
among Poloxamers,3-9 Triton X1009 and Tween 80.9 In some cases, nanospheres 
formation together with nanocapsules were observed. Aprotic, fully water-soluble 
solvents such as acetone and acetonitrile lead to high-quality nanocapsule preparations, 
whereas protic water-miscible solvents including ethanol, n-butanol, 
and isopropanol promoted the formation of nanospheres during nanocapsule 
preparation.5,9 It has been hypothesized that alcohols potentially initiate the polymerization 
reaction of alkylcyanoacrylates to form polymer nuclei or preformed 
polymers that may precipitate as nanospheres, when the organic phase is added to 
the acqueous phase.5 Lowering the pH in the organic phase was shown to inhibit 
polymerization in this medium.6 
Oil-containing nanocapsules have been used to encapsulate several types 
of biologically active compounds including both lipophilic molecules such as 
carbamazepine, indomethacin, lomustine, ethosuccimide, phenytoin,1,10-14 and 
hydrophilic drugs such as peptides.15-18 The lipophilic drugs were solubilized in the 
organic phase and were encapsulated during the preparation of the nanocapsules, 
usually using ethanol as the water-miscible organic solvent.4,17 The encapsulation 
efficiency of lipophilic drugs was found to be related to their solubility in the encapsulated 
oil.1 Quite surprisingly, hydrophilic compounds such as peptides have also 
been successfully encapsulated in oil-containing nanocapsules. Indeed, these highly 
water-soluble compounds do not tend to dissolve in oil. It has been suggested that 
the extremely rapid polymerization of the alkylcyanoacrylate occurring at the surface 
of the oil droplet limits the diffusion of the peptide towards the acqueous 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 259 
phase, therefore leading to its entrapment in nanocapsules.15 Another explanation 
is that surfactants may form inverse micelles in the oily phase, allowing some dissolution 
of hydrophilic compounds in this phase. Interestingly, in contrast to what 
has been observed with poly(alkylcyanoacrylate) nanospheres,19 peptides do not 
react chemically with the alkylcyanoacrylate monomer during the preparation of 
nanocapsules when ethanol is used. The presence of a large excess of alcohol seems 
to prevent the hydroxyl and amino groups of the peptides from reacting with the 
monomer, thus retaining the biological activity of the entrapped peptides.16-18,20'21 
For example, encapsulated insulin was still recognized by the insulin receptor of 
hepatocytes after nanoencapsulation.15,22 
2.1.2. Nanocapsules containing an acqueous core 
Nanocapsules with an acqueous core are a recent technology developed for the 
efficient encapsulation of water-soluble compounds, which are generally difficult 
to include within nanospheres. They were obtained by interfacial polymerization, 
where the alkylcyanoacrylates monomers were added to a W/O emulsion.23 
Anionic polymerization of the cyanoacrylate in the oily phase was initiated at the 
interface by nucleophiles such as hydroxyl ions in the acqueous phase, leading to 
the formation of nanocapsules with an acqueous core. In a typical procedure (Fig. 2), 
an acqueous phase at pH 7.4, consisted of ethanol and water, was prepared.23 This 
solution was emulsified in an organic phase containing Miglyol® and Montane® 80. 
The slow addition (4hrs) of the isobutylcyanoacrylate monomer in the organic 
phase under mechanical stirring allowed the polymerization to occur. This typical 
procedure leads to water droplets that are surrounded by a polymer core. The 
* " \ 9^ ' = Monomer 
CH2=CH 
COOR oo 
OILY PHASE 
Fig. 2. Schematic representation of the interfacial polymerization of cyanoacrylic 
monomers leading to the formation of nanocapsules with an acqueous core. 
260 Gref & Couvreur 
resuspension of the nanocapsules with a mean diameter approximately 350 nm in 
a water phase has been achieved by the ultracentrifugation of the oily suspension, 
with an excess of demineralized water containing a surfactant. After removal of the 
upper oily phase, the nanocapsules pellet was resuspended in water. 
These nanocapsules are very useful for the encapsulation of hydrophilic compounds 
such as oligonucleotides and peptides. In this case, these macromolecules 
are dissolved in the acqueous phase before the interfacial polymerization process 
takes place. For example, encapsulation efficiencies of 50% with an oligothymidylate 
(phosphodiester) and of 81% with a full phosphorothioate oligonucleotide 
(directed against EWS Fli-chimeric RNA) were obtained.23'24 These entrapment differences 
were attributed to possible interactions of the oligonucleotides with the 
oily phase, Montane® 80, or to the possible location of the oligonucleotide at the 
water-oil interface which could become saturated.24 
The localization of the oligonucleotide (within the acqueous core or adsorbed 
on the surface) has been investigated through fluorescence quenching experiments 
using fluorescein-labeled oligonucleotide and potassium iodine as an external 
quencher.23 It has been shown that fluorescent oligonucleotides were located in the 
acqueous core of the nanocapsules, surrounded by a polymeric wall, inaccessible to 
the quencher. On the contrary, when the fluorescent-oligonucleotides were free in 
solution, the fluorophores were highly accessible and strong quenching occurred. 
Similar quenching could be obtained with nanoencapsulated oligonucleotides only 
after the hydrolysis of the polymer wall, thus releasing the oligonucleotides. 
Zeta potential experiments have confirmed the localization of oligonucleotide 
in the acqueous core of the capsule.25 Moreover, nanoencapsulated oligonucleotides 
were protected against degradation by serum nucleases.25,26 Phosphorothioate 
oligonucleotides directed against EWS Fli-1 chimeric RNA encapsulated within 
poly(alkylcyanoacrylate) nanocapsules were tested in vivo for their efficacy against 
the experimental Ewing sarcoma in mice after intratumoral administration.24 Intratumoral 
injection of antisense-loaded nanocapsules led to a significant inhibition of 
tumor growth, whereas no antisense effect could be detected with the free oligonucleotide. 
These results were explained on the basis of a good protection of the 
oligonucleotide in the nanocapsules, which may act as a controled release system 
of oligonucleotide within the tumor. 
Salmon calcitonin was also successfully entrapped within poly (butylcyanoacrylate) 
nanocapsules of 300 nm in diameter.27 When the diameter was 
reduced to 50 nm, the encapsulation efficiencies decreased from 50 to 30%. After 
storage at room temperature or at 4°C, the nanocapsules retained their size for at 
least 34 months. The encapsulated calcitonin remained stable at 4°C for one year. 
Polyalkylcyanoacrylate nanocapsules were also prepared by interfacial polymerization, 
using a microemulsion instead of an emulsion as the template. 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 261 
Microemulsions are spontaneously forming, thermodynamically stable dispersed 
systems having a uniform droplet size of less than 200 nm. As such, they represent 
an interesting system that may be exploited for the preparation of nanocapsules 
too. Practically, a pseudo-ternary phase diagram of a mixture of medium 
chain glycerides (caprylic/capric triglycerides and mono-, diglycerides), a mixture 
of surfactants (polysorbate 80 and sorbitan monooleate) and water was constructed. 
Microemulsion domains were characterized by conductivity and viscosity 
to select systems suitable for the interfacial polymerization of ethyl-2-cyanoacrylate. 
Nanocapsules of 150 nm were obtained in those conditions and they were found to 
be able to encapsulate significant amounts of insulin.28 Size of the capsules may be 
controled, depending on different formulation variables.29 Factors influencing the 
encapsulation of hydrophilic compounds have been identified too.30 
2.2. Nanocapsules obtained from preformed polymers 
The preparation of nanocapsules from preformed polymers avoids some drawbacks 
of the interfacial polymerization process, such as the lack of control of the 
polymer molar masses and polydispersity, the presence of residual monomer in 
the preparation, and the possibility of drug inactivation.31 An interfacial deposition 
process to prepare nanocapsules, also known as nanoprecipitation, has been 
developed.32,33 In this simple and reproducible method, a water-miscible organic 
phase such as an alcohol or a ketone containing oil (with or without lipophilic 
surfactant) is mixed with an acqueous phase containing a hydrophilic surfactant. 
The preformed polymer, insoluble in both the oily and the acqueous phase, 
is solubilized in the organic phase. After the addition of the organic phase to 
the acqueous phase, the polymer diffuses with the organic solvent towards the 
acqueous phase and is stranded at the interface between oil and water. The driving 
force for nanocapsule formation is the rapid diffusion of the organic solvent 
in the acqueous phase, inducing interfacial nanoprecipitation of the polymer surrounding 
the droplets of the oily phase. Synthetic polymers such as poly(D,Llactide), 
poly(e-caprolactone) and poly(alkylcyanoacrylate) are most frequently 
employed for nanocapsule formation.32 Arabic gum, gelatin, ethylcellulose or 
hydroxypropylmethylcellulose phthalate were also successfully used.32 The size 
of nanocapsules is usually found between 100 and 500 nm, and it depends on 
several factors, namely, the chemical nature and the concentration of the polymer 
and the encapsulated drug, the amount of surfactants, the ratio of organic 
solvent to water, the concentration of oil in the organic solution, and the speed 
of diffusion of the organic phase in the acqueous phase. In general, the lower the 
interfacial tension and the viscosity of the oil, the smaller the nanocapsules are 
formed.34 
262 Gref & Couvreur 
Both lipophilic and hydrophilic surfactants are used in the preparation of 
nanocapsules by this technique. However, not all the surfactants that are technically 
suitable are acceptable for parenteral administration; as such, the choice 
has to be made with the administration route in mind. Generally, the lipophilic 
surfactant is a natural lecithin of relatively low phosphatidylcholine content, 
whereas the hydrophilic one is ionic (i.e. lauryl sulphate, quaternary ammonium), 
or more commonly nonionic (i.e. poly(oxyethylene)-poly(propropylene) 
glycol). 
Poly(ethylene glycol)-coated nanocapsules were also prepared by nanoprecipitation, 
using preformed diblock poly(lactide)-poly(ethylene glycol) copolymers 
or blends of these copolymers with the homopolymer poly(lactide.)35-38 
However, the most physically stable nanocapsules were those prepared with 
poly(lactide)-poly(ethylene glycol) copolymer alone. RU 58668, a promising pure 
antiestrogen, was entrapped into poly(ethylene glycol)-coated nanospheres and 
into nanocapsules with a similar coating.37 A series of preformed diblock polyesterpolyethylene 
glycol) copolymers were used for the design of these nanoparticles, 
both the molar masses of the poly(ethylene glycol) blocks and the nature 
of the hydrophobic polyester blocks being varied. Nanospheres which had a 
smaller size (~110nm), compared with nanocapsules (~250nm), were however 
able to incorporate larger amounts of the antioestrogen than the nanocapsules 
counterpart. 
In an alternative method named solvent displacement method, an O/W emulsion 
was formed.39 The organic phase contained the polymer, the oil and the drug, 
and the acqueous solution contained a stabilizing agent. In this procedure, the 
organic solvent was displaced into the external phase by the addition of an excess of 
water. This technique has several advantages such as the small quantities of solvents 
used, the good control of the size of the nanocapsules (80-900 nm), and the control 
of the thickness of the polymeric wall by monitoring the polymer concentrations.40 
However, large amounts of water have to be removed at the end of the 
process. 
Two formulation processes which bring lipids into play should also be mentioned. 
The first methodology is based on the inversion phase of an emulsion 
to prepare original lipidic nanocapsules. These capsules, interestingly obtained 
as a suspension in saline water, were constituted by medium chain triglycerides 
and hydrophilic /lipophilic surfactants. According to the authors, the formulation 
method has been developed to avoid the use of organic solvent or the high quantity 
of surfactants and co-surfactants, due to the potential toxicity of their residues after 
human administration. Their original structure was found to be a hybrid between 
polymeric nanocapsules and liposomes as their oily core is being surrounded by a 
tensioactive rigid membrane.41-43 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 263 
In another process, cisplatin lipid-based nanocapsules have been prepared by 
the repeated freezing and thawing of an equimolar dispersion of phosphatidylserine 
(PS) and phosphatidylcholine (PC) in a concentrated acqueous solution of 
cisplatin. Here, the molecular architecture of these novel nanostructures was elucidated 
by solid-state NMR techniques.15N NMR and 2H NMR spectra of nanocapsules 
containing 15N- and 2H-labeled cisplatin respectively, demonstrated that the 
core of the nanocapsules consists of solid cisplatin devoid of free water. Magicangle 
spinning 15N NMR showed that approximately 90% of the cisplatin in the 
core is present as the dichloro species. The remaining 10% was accounted for by 
a newly discovered dinuclear Pt compound that was identified as the positively 
charged chloride-bridged dimer of cisplatin. NMR techniques, sensitive to lipid 
organization 31P NMR and 2H NMR, revealed that the cisplatin core is coated by 
phospholipids in a bilayer configuration and that the interaction between solid 
core and bilayer coat exerts a strong ordering effect on the phospholipid molecules. 
Compared with phospholipids in liposomal membranes, the motion of the phospholipid 
headgroups is restricted and the ordering of the acyl chains is increased, 
particularly in PS.44 Analysis of the mechanism of the nanocapsule formation suggests 
that the method may be generalized to include other drugs showing low water 
solubility and lipophilicity.45 
3. Characterization 
Size evaluation of nanocapsules is most frequently done by photon correlation 
spectroscopy, transmission electron microscopy, and scanning electron microscopy, 
without or after freeze-fracture.33,39,46 At present, transmission electron microscopy 
performed after freeze-fracture has given the most useful information about 
nanocapsule structure, highlighting the polymer envelope and the inner cavity, 
and allowing the wall thickness to be estimated.1'7,47 Thus, polymer coatings were 
estimated to be around 5 ran, depending on the monomer concentration.47 Freezefracture 
(Fig. 1) has also allowed the visualization of different possible organizations 
of lipophilic surfactant, which can form vesicles, micelles, bilayers, or monolayers, 
depending on its concentration.33 The spherical shape of the nanocapsules was 
confirmed by atomic force microscopy.39 Most images of nanocapsules have been 
obtained by transmission electron microscopy performed on negatively stained 
preparations, allowing to gain information about nanocapsule morphology and 
integrity1,47 (Fig. 3A). Nanocapsules embedded in a suitable resin were cut into 
thin slices.48 They were observed using electron microscopy, the contrast being 
created by encapsulation of a colloidal gold-labeled molecule during nanocapsule 
preparation. In this manner, both polymer envelope and the internal cavity were 
distinguished easily (Fig. 3B). 
264 Gref& Couvreur 
50nm 100 nm 100 nm 
B 3 l 
Fig. 3. (A) Morphological appearance of polydactic acid-co-glycolic) nanocapsules using 
the transmission electron microscopy. (B) Labeling insulin with gold allows to distinguish the 
localization of this molecule into the internal core of poly(isobutyl cyanoacrylate) nanocapsules; 
Transmission Electron Microscopy. 
Zeta potential measurements are also very useful for the chraracterization of 
the nanocapsules. Surfactants and polymer are the major components that can affect 
this parameter. Many polymers such as poly (D,Llactide), poly(e-caprolactone) and 
lecithins impart a negative charge to the surface, whereas nonionic surfactants such 
as Poloxamer tend to reduce the absolute value of zeta potential.34 Calvo et alP 
described nanocapsules coated with positively charged polysaccharide chitosan. 
Their surface charge depended mainly on the viscosity of the chitosan solution used 
for coating. Positive values up to 46 mV were also observed with diethylaminoethyldextran 
coated nanocapsules.8 Generally, Zeta potential values above 30 mV (positive 
or negative values) lead to more stable nanocapsule suspensions, because 
repulsion between the particles prevented their aggregation. In contrast to observations 
with nanospheres, the negative Zeta potential of the nanocapsules was 
not completely masked by the presence of neutral poly(ethylene glycol) chains at 
the surface.63 This was due to the presence of lecithin in the polyethylene glycol) 
"brush", which remained necessary for nanocapsule stability. It was further highlighted 
that the presence of such a "brush" could reduce complement activation, 
an important step in the recognition of particles by macrophages.50'51 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 265 
Centrifugation in a density gradient was used to confirm the existence of 
nanocapsules by comparing with the colloidal carriers prepared without polymer 
or oil. For example, isopycnic centrifugation in a density gradient of Percoll 
was used in the case of nanocapsules with a Miglyol core and a coating of 
poly (alky lcyanoacry late) or poly(D,L lactide).39 The density of the nanocapsules 
was found to be intermediate between that of nanospheres and that of emulsions. 
These studies also demonstrated that the density of nanocapsules and the band 
thickness increased when the quantity of polymer increased. No contamination of 
nanocapsules with nanospheres was observed. However, Mosqueira et al.3i performed 
similar experiments and observed that nanocapsule preparations obtained 
by nanoprecipitation contained small amounts of nanospheres, as it has previously 
been described by Gallardo et al.5 for nanocapsules prepared by interfacial 
polymerization. When lecithin was present in excess as lipophilic surfactant, liposomes 
were also detected in the nanocapsule preparations. Liposomes could not 
be distinguished from nanocapsules on the basis of density differences, but have 
been detected by electron microscopy52 and by the encapsulation of an acqueous 
tracer.34 
4. Drug Release 
Release of encapsulated drugs from nanocapsules made of preformed polymers, 
appears only to be controled by the partition coefficient of the drug between the 
oily core and the acqueous external medium, and the relative volumes of these 
two phases. Except for macromolecules, the rate of diffusion of the drug through 
the thin polymeric coating does not seem to be a limiting factor, nor does the 
nature of the polymeric wall. This clearly suggests that the polymer membrane 
may be porous rather than a continuous film barrier to diffusional release. The 
nature of the external acqueous phase is of prime importance in the release. For 
example, indomethacin release was faster and more complete in the presence of 
albumin, which acts as an acceptor in the acqueous phase.11,52 Similarly, release 
of halofantrine, a highly lipophilic drug, was only observed in the presence of 
serum, because the drug has a high affinity for lipoproteins.36 The presence of 
a hydrophilic poly(ethylene glycol) "brush" at the nanocapsule surface was also 
shown to play a role in drug release. Release of halofantrine and primaquine from 
such surface-modified nanocapsules was reduced, compared with conventional 
nanocapsules.36,53 
In conclusion, it may be considered a challenge to develop nanocapsule systems 
with release profiles, which may be controled not only by the partitioning 
coefficient, but also by the nature or morphology (i.e. thickness or porosity) of the 
surrounding membrane. 
266 Gref & Couvreur 
5. Applications 
Nanocapsules have been proposed as drug delivery systems for several drugs by 
different routes of administration such as oral, ocular or parenteral. Drug-loaded 
nanocapsules were used to improve the stability of the drug either in biological 
fluids, or simply in the formulation. Another goal was to reduce the toxicity of 
some drugs known for their undesirable side effects. 
5.1. Oral route 
Challenging aspects related to oral administration deal with the entrapment of 
unstable molecules, such as peptides or that of anti-inflammatory compounds that 
cause local side effects on the mucosae. Pioneering studies in the mid 1980s dealt 
with indomethacin and insulin entrapment. 
Indomethacin, an anti inflammatory drug, has been successfully encapsulated 
in the polyalkylcyanoacrylate nanocapsules with the aim of reducing its side effects 
on the gastric and intestinal mucosa.11 The drug retained its biological activity after 
nanoencapsulation. Moreover, nanoencapsulated formulations allowed a dramatic 
reduction of the ulcerative side effects usually induced by indomethacin on the 
mucosae.54 This protection was attributed to the combined effect of the sustained 
release of indomethacin from the nanocapsules, with a significant reduction of the 
direct contact between drug and the mucosae. In the case of nanocapsules obtained 
by nanoprecipitation using polyesters, the release kinetics in media mimicking 
pH of the gut were more sensitive to changes in drug partitioning related to the 
change of pH, than to the type of polymer used.55,56 Drug release from nanocapsules 
was accelerated in the presence of digestive enzymes such as proteases and 
esterases. This was correlated with a decrease in polymer molecular weight.55'56 
Diclofenac and indomethacin, two major nonsteroidal anti inflammatory agents, 
have been encapsulated in polyQactic acid) nanocapsules obtained by nanoprecipitation, 
with the aim of reducing their side effects on the gastric mucosa.54,57,58 The 
side effects of both drugs were completely modified and reduced by the encapsulation 
in nanocapsules.54 As in the case of nanocapsules produced by interfacial 
polymerization, a marked protective effect on the gastrointestinal mucosa, as compared 
with the ulcerative effect observed with the drug solutions, was observed. 
Insulin-loaded nanocapsules yielded promising pharmacological results.16,21 
When given orally to diabetic rats and dogs, single administration produced 
a reduction in glycemia after an unusually long lag of several days, and this 
hypoglycemia was sustained for up to 20 days.16,20,21,59 It was suggested that 
nanocapsules could release insulin slowly from a depot within the body. The 
nanocapsules seemed to be involved in carrying the insulin near the intestinal 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 267 
epithelium where they were absorbed and translocated as intact nanocapsules to the 
blood vessels.48,59-61 However, Lowe and Temple16 reported that insulin adsorption 
from orally administered nanocapsules reached a maximum of absorption, 15 min 
after administration and any trace of insulin in blood was detected after a few 
hours. Sai et al.62 have proposed the use of insulin-loaded nanocapsules as a new 
prophylactic tool to prevent diabetes. They showed in a model of non-obese diabetic 
mice that prophylactic injection of such nanocapsules reduced the incidence of 
diabetes. 
Anti infectious agents such as atovaquone and rifabutin, two compounds active 
against the opportunistic parasite Toxoplasma gondii, were successfully entrapped in 
poly(lactide) nanocapsules formed by nanoprecipitation. These drugs have a poor 
bioavailability because of their insolubility in water. Nanoencapsulation is allowed 
to decrease in the brain parasitic burden in a higher extent than the same amount 
of free drug.63 
Chitosan-coated nanocapsules were particularly interesting for oral administration, 
probably because their positive charge allow them to stick efficiently along 
the gastro-intestinal mucosa, with a further possible diffusion through the epithelium 
, thus providing a continuous drug delivery into the blood stream.64,65 When 
the peptide salmon calcitonin was entrapped into these nanocapsules, long-lasting 
hypocalcemia effects were observed, following oral administration to rats.66 In contrast, 
calcitonin control emulsions led to negligible responses. 
5.2. Parenteral route 
As far as the parenteral route is concerned, nanocapsules could be useful for the 
formulation of poorly soluble drugs, and for controling the drug biodistribution 
according to the properties of the carrier. In this view, indomethacin and diclofenac 
were entrapped in nanocapsules, but diclofenac in solution or in nanocapsules 
showed similar plasma concentration profiles. After intravenous administration, 
encapsulated indomethacin showed even lower plasma concentrations than the 
free drug because of enhanced hepatic uptake of loaded nanocapsules.57 One possible 
explanation for the absence of the modification of the pharmacokinetics and 
biodistribution profiles of the encapsulated drugs probably results from the rapid 
rate of release of these drugs into the circulation, due to the high blood dilution 
and/or the presence of plasma proteins. Subcutaneous injection did not lead to a 
slow release of the drug either. Nevertheless, after intramuscular administration, 
the nanocapsules containing diclofenac showed a significantly reduced inflammation 
at the site of injection, compared with the free drug in solution.67 Similarly, 
darodipine nanocapsules provided a prolonged antihypertensive effect compared 
with free drug which lasted for at least 24 hrs.68 
268 Gref & Couvreur 
Nanocapsules prepared by interf acial polymerization of the isobutylcyanoacrylate 
monomers were retained longer at the injection site after intramuscular administration 
than the other types of carriers such as emulsions or liposomes.69 Moreover, 
they were taken up to a significant extent by the regional lymph nodes, likely owing 
to the phagocytosis by macrophages. These observations open up the possibility of 
delivering cytostatic drugs and immunomodulators to the lymph node metastases. 
When administered intravenously, nanocapsules made by interfacial polymerization 
or by nanoprecipitation were taken up rapidly by organs of the mononuclear 
phagocyte system, mainly the liver.70 To take advantage of this particular tissue 
distribution, nanocapsules containing muramyltripeptide cholesterol (MTPChol) 
were designed.71,72 This immunostimulating agent, able to activate the 
macrophages and to stimulate their innate defense functions against tumor cells, is a 
useful agent to treat metastatic cancer. In vitro studies with rat alveolar macrophages 
have shown that nanocapsules prepared from poly(D,Llactic acid) containing MTPChol 
were more efficient activators than the free drug. This was attributed to the 
intracellular delivery of the nanoencapsulated immunomodulator after cell phagocytosis; 
an intermediate transfer of the drug to serum proteins was another suggested 
mechanism.73 In vivo, this type of nanocapsules is allowed to obtain significant 
antimetastatic effects in a model of liver metastases.74 
For other types of applications, to avoid the rapid clearance by the mononuclear 
phagocyte system, nanocapsules coated with poly(ethylene glycol) with a 
molar mass of 20,000 g/mole were developed. An antimalarial drug, halofantrine, 
was entrapped with the aim of obtaining a well-tolerated injectable form for the 
treatment of this severe intravascular disease.36 In mice, at an advanced stage of 
infection with Plasmodium berghei, the area under the curve for plasma halofantrine 
was increased six-fold, compared with the free drug when the molecule was presented 
as nanocapsules. Moreover, the toxicity of halofantrine was reduced by 
incorporation into the nanocapsules. Up to 100 mg/kg could be administered intravenously 
without toxicity, yet all mice injected with this dose of free halofantrine 
died instantaneously. However, in vivo, only small differences were observed in 
terms of the therapeutic activity between poly(ethylene glycol) coated nanocapsules 
and the uncoated ones. This was explained by the possible saturation of the 
phagocytic capacity of the liver in severely infected mice, as a result of the uptake of 
parasitized erythrocytes.75 Moreover, it was emphasized that the amount of serum 
lipoproteins, which acted as acceptors for halofantrine released from nanocapsules, 
is reduced during the disease. 
Poly(ethylene glycol) coated nanocapsules were also used to deliver lipophilic 
drugs to the solid tumors. In this case, the vascular endothelium is known to be 
more permeable, thus allowing the extravasation of small-sized colloidal particles. 
This specific distribution of colloids into tumoral sites is known as the enhanced 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 269 
permeability and retention effect (EPR effect). The efficacy of this strategy has been 
demonstrated using a photosensitizer, meta-tetra (hydroxyphenyl) chlorine, encapsulated 
in nanocapsules designed from diblock poly(D,L lactide)-poly(ethylene glycol) 
copolymers.35 
5.3. Ocular delivery 
The major problems encountered when delivering drugs to the eyes are the poor 
permeability of the corneal epithelium and the rapid clearance because of tear 
turnover and lacrimal drainage. Nanocapsule formulations were developed with 
the aim of improving drug efficacy by retaining it at the level of the ocular tissue, 
thus reducing the number of administrations.7677 
Betaxolol-loaded poly(isobutylcyanoacrylate) nanocapsules made by interracial 
polymerization were prepared for the treatment of glaucoma. Only a marginal 
decrease in the intraocular pressure was observed with this type of formulation, 
compared with the activity obtained with the commercial form (single solution) or 
by other carriers.13 More promising results have been obtained with pilocarpine.14 
In this case, sustained drug release was obtained when incorporating the pilocarpine 
loaded nanocapsules into a Pluronic gel. Thus, a significant increase in the 
bioavailability of the drug was achieved. 
Ganciclovir is an antiviral drug used for the treatment of cytomegalovirus infections. 
In the clinical practice, two to three intravitreal injections per week are needed 
to overcome the rapid clearance of the drug from the eyes. Ganciclovir encapsulation 
in poly(ethylcyanoacrylate) nanocapsules made by interfacial polymerization 
provided a sustained release of the drug over four days.10 Moreover, after intravitreal 
injection of the nanocapsules, the drug could still be detected in the eyes at a 
therapeutic level after ten days. Significant amounts of ganciclovir were found in 
the retina and in the vitreous humor which is considered as beneficial in the treatment 
of cytomegalovirus retinitis. On the contrary, after administration of single 
solutions of the drug free, the maximum concentration of ganciclovir was reached 
in less than one day and no drug could be detected later. However, despite these 
beneficial results, some toxicity (opacification of the lens and vitreous humor turbidity) 
was found as a result of the nanocapsules. 
Antiglaucomatous agents such as carteolol and betaxolol were also encapsulated 
in nanocapsules prepared from preformed polymers, but they only showed 
a reduction of the noncorneal absorption (systemic circulation), leading to lesser 
side effects as compared with the free drug.13'78'79 Encapsulation in nanocapsules 
produced an improved pharmacological effect characterized by a more important 
reduction of the intraocular pressure, compared with the free drug treatment, 
as well as with the same treatment but delivered by nanospheres; reduced 
270 Gref & Couvreur 
cardiovascular systemic side effects were also observed with the nanocapsules. ' 
In the case of betaxolol, the nature of the polymer making up the nanocapsule 
wall was found to play a major role in the pharmacological responses.78'80 Thus, 
poly(e-caprolactone) walls were more efficient than poly(isobutylcyanocrylate) 
or poly(lactide-co-glycolide) ones. Indeed, as shown by the confocal microscopy, 
poly(e-caprolactone) nanocapsules could specifically penetrate the corneal epithelium 
by an endocytic process, without causing any damage to the cells. In contrast, 
poly(isobutylcyanoacrylate) nanoparticles produced a cellular lysis.81 As no differences 
in penetration were observed between nanospheres and nanocapsules, 
the presence of an oily core did not seem to influence activity of the formulation. 
Coating the negatively charged surface of poly(e-caprolactone) nanocapsules with 
chitosan, a cationic polymer, provided the best corneal drug penetration, together 
with preventing the degradation caused by the adsorption of lysozyme, a positively 
charged enzyme found in tear fluid.82 This was explained by the higher penetration 
of the nanocapsules into the corneal epithelial cells and by the mucoadhesion of 
these positively charged particles onto the negatively charged membranes. Additionally, 
a specific effect of chitosan on the tight junctions has been mentioned.83 
Encouraging results were also obtained with nanocapsules containing the 
immunosuppressive peptide cyclosporin A.84 This drug was efficiently entrapped 
in poly(e-caprolactone) nanocapsules, leading to a five-fold increase of the 
cyclosporin A corneal concentrations, compared with an oily solution of the drug. 
Again, chitosan-overcoated nanocapsules were able to provide a selective and prolonged 
delivery of cyclosporine A to the ocular mucosae, without compromising 
the inner ocular tissues and avoiding systemic absorption.84 The mechanism that 
explains the increased ocular penetration was understood as the combination of an 
improved interaction with the corneal epithelium, followed by the penetration of 
the particles into the corneal epithelium.85 In the case of indomethacin associated 
with chitosane-coated nanocapsules, the use of confocal microscopy established the 
fact that the nanocapsules penetrated through the corneal epithelium following a 
transcellular pathway.85,86 
6. Conclusion 
As discussed in this chapter, there are now various technologies for the preparation 
of nanocapsules. These methods which obey a wide variety of principles may either 
start from a monomer or from a preformed polymer. They employ macromolecular 
materials of synthetic or natural origin and they allow the design of nanocapsules 
with either an acqueous or an oily core. Thus, they can efficiently entrap almost 
every molecule. The most significant advantage of nanocapsules over nanospheres 
is that the drug to polymer ratio is generally much higher, which allows the use of 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 271 
lesser polymer to deliver the same amount of drug to the cells and tissues. This is, 
from a toxicological point of view, a substantial advantage of this type of technology. 
On the contrary, drug release from nanocapsules is mainly dependent on the partitioning 
coefficient of the biologically active compound between the nanocapsule 
core and the biological receptor medium. If the nanocapsule thin polymer membrane 
may be a barrier for the diffusion of macromolecules, it is not the case for 
small organic molecules. Thus, to control the drug release kinetic from nanocapsules, 
it is likely to remain the primary challenge to be resolved with this kind of 
technology in the next few years. 
References 
1. Fresta M, Cavallaro G, Giammona G, Wehrli E and Puglisi G (1996) Preparation and characterization 
of polyethyl-2-cyanoacrylate nanocapsules containing antiepileptic drugs. 
Biomaterials 17:751. 
2. Fattal E, Peracchia MT and Couvreur P (1997) Poly(alkylcyanoacrylate), in: Domb AJ, 
Kost J, Wiseman DM (eds.) Handbook of Biodegradable Polymers. Harwood Academic 
Publishers: Amsterdam. 
3. Al Khouri N, Fessi H, Roblot-Treupel L, Devissaguet JP and Puisieux F (1986) An original 
procedure for preparing nanocapsules of polyalkyl-cyanoacrylates for interfacial 
polymerization. Pharm Acta Helv 61:274. 
4. Al Khouri Fallouh N, Roblot-Treupel L, Fessi H, Devissaguet JP and Puisieux F (1986) 
Development of a new process for the manufacture of polyisobutylcyanoacrylate 
nanocapsules. Int J Pharm 28:125. 
5. Gallardo MM, Couarraze G, Denizot B, Treupel L, Couvreur P and Puisieux F (1993) 
Preparation and purification of isohexylcyanoacrylate nanocapsules. Int ] Pharm 100:55. 
6. Wohlgemuth M, Machtle W and Mayer C (2000) Improved preparation and physical 
studies of polybutylcyanoacrylate nanocapsules. / Microencapsulation 17:437. 
7. Chouinard F, Kan FW, Leroux JC, Foucher C and Lenaerts V (1991) Preparation and 
purification of polyisohexylcyanoacrylate nanocapsules. Int} Pharm 72:211. 
8. Chouinard F, Buczkowski S and Lenaerts V (1994) Poly(alkylcyanoacrylate) nanocapsules: 
Physicochemical characterization and mechanism of formation. Pharm Res 11:869. 
9. Puglisi G, Fresta M, Giammona G and Venture CA (1995) Influence of the preparation 
conditions in poly(ethylcyanoacrylate) nanocapsules formation. Int J Pharm 125:283. 
10. El-Samaligy MS, Rojanasakul Y, Charlton JF, Weinstein GW and Lim JK (1996) Ocular 
disposition of nanoencapsulated acyclovir and ganciclovir via intravitreal injection in 
rabbit's eye. Drug Del 3:93. 
11. Ammoury N, Fessi H, Devissaguet JP, Puisieux F and Benita S (1989) Physicochemical 
characterization of polymeric nanocapsules and in vitro release evaluation of 
indomethacin as a drug model. STP Pharm 5:642. 
12. Giirsoy A, Eroglu L, Ulutin S, Tasyiirek M, Fessi H, Puisieux F and Devissaguet JP (1989) 
Evaluation of indomethacin nanocapsules for their physical stability and inhibitory 
activity on inflammation and platelet aggregation. Int J Pharm 52:101. 
272 Gref& Couvreur 
13. Marchal-Heussler L, Fessi H, Devissaguet JP, Hoffman M and Maincent P (1992) 
Colloidal drug delivery systems for the eye. A comparison of the efficacy of three different 
polymers: Polyisobutylcyanoacrylate, polylactic-co-glycolic acid, poly-epsiloncaprolactone. 
STP Pharm Sci 2:98. 
14. Desai SD and Blanchard J (2000) Pluronic® F127-based ocular delivery systems containing 
biodegradable polyisobutylcyanoacrylate nanocapsules of pilocarpine. Drug Del J 
7:201. 
15. Aboubakar M, Puisieux F, Couvreur P, Deyme M and Vauthier C (1999) Study of 
the mechanism of insulin encapsulation in poly(isobutylcyanoacrylate) nanocapsules 
obtained by interfacial polymerization. / Biomed Mater Res 47:568. 
16. Damge C, Michel C, Aprahamiam M and Couvreur P (1988) New approach for oral 
administration of insulin with polyalkylcyanoacrylate nanocapsules as drug carrier. 
Diabetes 37:246. 
17. Lowe PJ and Temple CS (1994) Calcitonin and insulin in isobutylcyanoacrylate nanocapsules: 
Protection against proteases and effect in intestinal absorption in rats. / Pharm 
Pharmacol 46:547. 
18. Damge C, Vonderscher J, Marbach P and Pinget M (1997) Poly(alkylcyanoacrylate) 
nanocapsules as a delivery system in the rat for octreotide, a long-acting somatostatin 
analogue. / Pharm Pharmacol 49:949. 
19. Damge C, Vranckx H, Baldschmidt P and Couvreur P (1997) Poly(alkylcyanoacrylate) 
nanospheres for oral administration of insulin. / Pharm Sci 86:1403. 
20. Michel C, Aprahamiam M, Defontaine L, Couvreur P and Damge C (1991) The effect 
of site of administration in the gastrointestinal tract on the absorption of insulin from 
nanocapsules in diabetic rats. / Pharm Pharmacol 43:1. 
21. Damge C, Hillaire-Buys D, Puech R, Hoelizel A, Michel C and Ribes G (1995) Effects 
of orally administered insulin nanocapsules to normal and diabetic dogs. Diabetes Nutr 
Metab 8:3. 
22. Roques M, Damge C, Michel C, Staedel C, Cremel G and Hubert P (1992) Encapsulation 
of insulin for oral administration preserves interaction of the hormone with its receptor 
in vitro. Diabetes 41:451. 
23. Lambert G, Fattal E, Pinto-Alphandary H, Gulik A and Couvreur P (2000) Polyisobutylcyanoacrylate 
nanocapsules containing an acqueous core as a novel colloidal carrier for 
the delivery of oligonucleotides. Pharm Res 17:707. 
24. Lambert G, Bertrand JR, Fattal E, Subra F, Pinto-Alphandary H, Malvy C, Auclair C and 
Couvreur P (2000) EWS Fli-1 antisense nanocapsules inhibits Ewing sarcoma-related 
tumor in mice. BBRC 279:401. 
25. Chavany C, Le Doan T, Couvreur P, Puisieux F and Helene C (1992) Polyalkylcyanoacrylate 
nanoparticles as polymeric carriers for antisense oligonucleotides. Pharm Res 
9:441. 
26. Nakada Y, Fattal E, Foulquier M and Couvreur P (1996) Pharmacokinetics and biodistribution 
of oligonucleotide adsorbed onto poly(isobutylcyanoacrylate) nanoparticles 
after intravenous administration in mice. Pharm Res 13:38. 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 273 
27. Vrankx H, Demoustier M and Deleers M (1996) A new nanocapsule formulation with 
hydrophilic core: Application to the oral administration of salmon calcitonin in rats. Eur 
J Pharm Biopharm 42:345. 
28. Watnasirichaikul S, Davies NM, Rades T and Tucker IG (2000) Preparation of biodegradable 
insulin nanocapsules from biocompatible microemulsions. Pharm Res 17:684-689. 
29. Watnasirichaikul S, Rades T, Tucker IG and Davies NM (2002) Effects of formulation 
variables on characteristics of poly (ethylcyanoacrylate) nanocapsules prepared from 
w/o microemulsions. Int J Pharm 235:237-246. 
30. Pitaksuteepong T, Davies NM, Tucker IG and Rades T (2002) Factors influencing the 
entrapment of hydrophilic compounds in nanocapsules prepared by interfacial polymerisation 
of water-in-oil microemulsions. Eur } Pharm Biopharm 53:335-342. 
31. Grangier JL, Puygrenier M, Gauthier JC and Couvreur P (1991) Nanoparticles as carriers 
for growth hormone releasing factor (GRF). / Control Rel 15:3. 
32. Fessi H, Devissaguet JP and Puisieux F (1986) Procede de preparation des systemes 
collo'idaux dispersibles d'une substance sous forme de nanocapsules. French Patent 
Application No. 8618444. 
33. Fessi H, Puisieux F, Devissaguet JP, Ammoury N and Benita S (1989) Nanocapsule formation 
by interfacial deposition following solvent displacement. Int J Pharm 55:R1-R4. 
34. Mosqueira VCF, Legrand P, Pinto-Alphandary H, Puisieux F and Barratt G (2000) 
Poly(D,L-lactide) nanocapsules prepared by a solvent displacement process: Influence 
of the composition on physicochemical and structural properties. / Pharm Sci 89:614. 
35. Bourdon O, Mosqueira V, Legrand P and Blais J (2000) A comparative study of the 
cellular uptake, localization and phototoxicity of meta-tetra(hydroxyphenyl) chlorine 
encapsulated in surfacemodified submicronic oil/water carriers in HT29 tumor cells. 
/ Photochem Photobiol B 55:164. 
36. Mosqueira VCF, Legrand P, Gref R and Barratt G (1999) In vitro release kinetic studies 
of PEG-modified nanocapsules and nanospheres loaded with a lipophilic drug: Halofantrine 
base. Proceedings of the 26th International Symposium on Controlled Release ofBioactive 
Materials, 20-23 June, Boston, MA, 1074-1075. 
37. Ameller T, Marsaud V, Legrand P, Gref R and Barratt G et Renoir JM (2003) Polyester — 
poly(ethylene glycol) nanospheres and nanocapsules loaded with the pure antioestrogen 
RU 58668: Physico-chemical and opsonisation properties. Pharm Res 20(7):1063-1070. 
38. Ameller T, Marsaud V, Legrand P, Gref R and Barratt G et Renoir JM (2003) Polyester — 
poly(ethylene glycol) nanospheres and nanocapsules loaded with the pure antioestrogen 
RU 58668: Physico-chemical and opsonisation properties. Pharm Res 20(7):1063-1070. 
39. Quintanar-Guerrero D, Allemann E, Doelker E and Fessi H (1998) Preparation and 
characterization of nanocapsules from preformed polymers by a new process based 
on emulsification-diffusion technique. Pharm Res 15:1056. 
40. Quintanar-Guerrero D, Fessi H, Doelker E and Allemann E (1997) Procede de preparation 
de nanocapsules de type vesiculate utilisable notamment comme vecteurs collo'idaux 
de principes actifs pharmaceutiques ou autres, French Patent 97.09.672. 
41. Heurtault B, Saulnier P, Pech B, Proust JE and Benoit JP (2002) A novel phase inversionbased 
process for the preparation of lipid nanocarriers. Pharm Res 19:875-880. 
274 Gref & Couvreur 
42. Heurtault B, Saulnier P, Pech B, Proust JE and Benoit JP (2002) Properties of 
polyethyleneglycol 660 12 hydroxystearate at the triglyceride/water interface. Int J 
Pharm 242:176-170. 
43. Heurtault B, Saulnier P, Pech B, Venier-Julienne MC, Proust JE, Phan-Tan-Luu R and 
Benoit JP (2003) The influence of lipid nanocapsule composition on their size distribution. 
Eur Pharm Sci 18:55-61. 
44. Chupin V, de Kroon AI and de Kruijff B (2004) Molecular architecture of nanocapsules, 
bilayer-enclosed solid particles of Cisplatin. / Am Chem Soc 126:13816-13821. 
45. Burger KN, Staffhorst RW, de Vijlder HC, Velinova MJ, Bomans PH, Frederik PM and 
de Kruijff B (2002) Nanocapsules: Lipid-coated aggregates of cisplatin with high cytotoxicity. 
Nat Med 81-84. 
46. Magalhaes NS, Fessi H, Puisieux F, Benita S and Seiller M (1995) An in vitro release kinetic 
examination and comparative evaluation between submicron emulsion and polylactic 
acid nanocapsules of clofibride. / Microencapsul 12:195. 
47. Rollot JM, Couvreur P, Roblot-Treupel L and Puisieux F (1986) Physicochemical and 
morphological characterization of polyisobutyl cyanoacrylate nanocapsules. / Pharm 
Sci 75:361. 
48. Vauthier C, Dubernet C, Fattal E, Pinto-Alphandary H and Couvreur P (2003) 
Poly(alkylcyanoacrylates) as biodegradable materials for biomedical applications. Adv 
Drug Del Rev 55(4):519-548. 
49. Calvo P, Remunan-Lopez C, Vila-Jato JL and Alonso MJ (1997) Development of positively 
charged colloidal drug carriers: Chitosan-coated polyester nanocapsules and submicronemulsions. 
Coll Polym Sci 275:46. 
50. Mosqueira VCF, Legrand P, Gulik A, Bourdon O, Gref R, Labarre D and Barratt G (2001) 
Relationship between complement activation, cellular uptake and surface physicochemical 
aspects of novel PEG-modified nanocapsules. Biomaterials 22:2967. 
51. Mosqueira VCF, Legrand P, Gref R, Heurtault B, Appel M and Barratt G (1999) Interactions 
between a macrophage cell line (J774A1) and surface-modified poly(D,L-lactide) 
nanocapsules bearing poly(ethylene glycol). / Drug Targ 7:65. 
52. Ammoury N, Fessi H, Devissaguet JP, Puisieux F and Benita S (1990) In vitro release 
kinetic pattern of indomethacin from poly(D,L-lactide) nanocapsules. / Pharm Sci 79:763. 
53. Heurtault B, Legrand P, Mosqueira V, Devissaguet JP, Barratt G and Bories C (2001) 
In vitro antileishmanial properties of surface modified and primaquine loaded nanocapsules 
in Leishmania donovani intramacrophagic amastigotes. Ann Trop Med Parasitol 
95:529. 
54. Ammoury N, Fessi H, Devissaguet JP, Dubrasquet M and Benita S (1991) Jejunal absorption, 
pharmacological activity, and pharmacokinetic evaluation of indomethacin-loaded 
poly(D,L-lactide) and poly(isobutyl-cyanoacrylate) nanocapsules in rats. Pharm Res 
8:101. 
55. Marchais H, Benali S, Irache JM, Tharasse-Bloch C, Lafont O and Orecchioni AM (1998) 
Entrapment efficiency and initial release of phenylbutazone from nanocapsules prepared 
from different polyesters. Drug Dev Ind Pharm 24:883. 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 275 
56. Kedzierewicz F, Thouvenot P, Monot I, Hoffman M and Maincent P (1998) Influence of 
different physicochemical conditions on the release of indium oxine from nanocapsules. 
J Biomed Mater Res 39:588. 
57. Guterres SS, Fessi H, Barratt G, Puisieux F and Devissaguet JP (1995) Poly(D,L-lactide) 
nanocapsules containing non-steroidal anti-inflammatory drugs: Gastrointestinal tolerance 
following intravenous and oral administration. Pharm Res 12:1545. 
58. Andrieu V, Fessi H, Dubrasquet M, Devissaguet JP, Puisieux F and Benita S (1989) Pharmacokinetic 
evaluation of indomethacin nanocapsules. Drug Des Del 4:295. 
59. Damge C, Michel C, Aprahamiam M, Couvreur P and Devissaguet JP (1990) Nanocapsules 
as carriers for oral peptide delivery. / Control Rel 13:233. 
60. Aboubakar M, Couvreur P, Pinto-Alphandary H, Gouritin B, Lacour B, Farinotti R, et al. 
(2000) Insulin-loaded nanocapsules for oral administration: In vitro and in vivo investigation. 
Drug Dev Res 49:109. 
61. Aprahamiam M, Michel C, Humbert W, Devissaguet JP and Damge C (1987) Transmucosal 
passage of polyalkylcyanoacrylate nanocapsules as a new drug carrier in the small 
intestine. Biol Cell 61:69. 
62. Sai P, Damge C, Rivereau AS, Hoeltzel A and Gouin E (1996) Prophylactic oral administration 
of metabolically active insulin entrapped in isobutylcyanoacrylate nanocapsules 
reduces the incidence of diabetes in nonobese diabetic mice. ] Autoimmunity 9:713. 
63. Dalengon F, Amjaud Y, Lafforgue C, Derouin F and Fessi H (1998) Atovaquone and 
rifabutine-loaded nanocapsules: Formulation studies, hit J Pharm 153:127. 
64. Calvo P, Remufian C, Vila Jato JL and Alonso MJ (1997) Development of positively 
charged colloidal drug carriers: Chitosan-coated polyester nanocapsules and submicron 
emulsions. Coll Polym Sci 275:46-53 
65. Vila A, Sanchez A, Evora C, Soriano I, Vila Jato JL and Alonso MJ (2004) PEG-PLA 
nanoparticles as carriers for nasal protein/vaccine delivery. / Aerosol Med. 
66. Prego C, Fernandez-Megia E, Novoa-Carballal R, Quinoa E, Torres D and Alonso MJ 
(2003) Chitosan and Chitosan-PEG nanocapsules: New carriers for improving the oral 
absorption of calcitonin. 
67. Guterres SS, Fessi H, Barratt G, Puisieux F and Devissaguet JP (2000) Poly(D,L-lactide) 
nanocapsules containing diclofenac: Protection from muscular damage in rats. / Biomater 
Sci Polym Ed 11:1347. 
68. Hubert B, Atkinson J, Guerret M, Hoffman M, Devissaguet JP and Maincent P (1991) The 
preparation and acute antihypertensive effects of a nanocapsular form of darodipine, a 
dihydropyridine calcium entry blocker. Pharm Res 8:734. 
69. Nishioka Y and Yoshino H (2001) Lymphatic targeting with nanoparticulate system. Adv 
Drug Del Rev 47:55. 
70. Marchal-Heussler L, Thouvenot P, Hoffman M and Maincent P (1999) Comparison of the 
biodistribution in mice of Ill-indium oxine encapsulated into poly(lactic-co-glycolic)- 
D,L-85/15 and poly(epsilon-caprolactone) nanocapsules. ] Pharm Sci 88:450. 
71. Morin C, Barratt G, Fessi H, Devissaguet JP and Puisieux F (1994) Improved intracellular 
delivery of a muramyl dipeptide analog by means of nanocapsules. Int J Immunopharmacol 
16:461. 
276 Gref& Couvreur 
72. Seyler I, Appel M, Devissaguet JP Legrand P and Barratt G (1996) Relationship between 
NO-synthase activity and TNF-alpha secretion in mouse macrophage lines stimulated 
by a muramyl peptide entrapped in nanocapsules. Int J Immunopharmacol 18:385. 
73. Seyler I, Appel M, Devissaguet JP, Legrand P and Barratt G (1999) Macrophage activation 
by a lipophilic derivative of muramyldipeptide within nanocapsules: Investigation of 
the mechanism of drug delivery. / Nanoparticle Res 1:91. 
74. Barratt G, Puisieux F, Yu WP, Foucher C, Fessi H and Devissaguet JP (1994) Antimetastatic 
activity of MDP-L-alanyl-cholesterol incorporated into various types of 
nanocapsules. Int J Immunopharmacol 16:457. 
75. Mosqueira VCF, Legrand P, Bories C, Devissaguet JP and Barratt G (2000) Comparative 
pharmacokinetics and in vivo efficacy of an intravenous formulation of halofantrine in 
long-circulating nanocapsules in Plasmodium berghei-iniected mice. Proceedings of the 27th 
International Symposium on Controlled Release of Bioactive Materials, 7-13 July Paris, France; 
490-491. 
76. Langer K, Zimmer A and Kreuter J (1999) Acrylic nanoparticles for ocular drug delivery. 
STP Pharm Sci 7:445. 
77. Legrand P, Barratt G, Mosqueira V, Fessi H and Devissaguet JP (1999) Polymeric nanocapsules 
as drug delivery systems, a review. STP Pharm Sci 9:411. 
78. Marchal-Heussler L, Sirbat D, Hoffman M and Maincent P (1993) Poly(epsiloncaprolactone) 
nanocapsules in carteolol ophthalmic delivery. Pharm Res 10:386. 
79. Losa C, Marchal-Heussler L, Orallo F, Vila Jato JL and Alonso MJ (1993) Design of 
new formulations for topical ocular administration: Polymeric nanocapsules containing 
metipranolol. Pharm Res 10:80. 
80. Calvo P, Thomas C, Alonso MJ, Vila Jato JL and Robinson JR (1994) Study of the mechanism 
of interaction of poly(e-caprolactone) nanocapsules with the cornea by confocal 
laser scanning microscopy. Int} Pharm 103:283. 
81. Zimmer A, Kreuter J and Robinson JR (1991) Studies on the transport pathway of PBCA 
nanoparticles in ocular tissues. / Microencapsulation 8:497. 
82. Calvo P, Vila-Jato JL and Alonso MJ (1997) Effect of lysozyme on the stability of polyester 
nanocapsules and nanoparticles: Stabilization approaches. Biomaterials 18:1305. 
83. Calvo P, Vila-Jato JL and Alonso MJ (1997) Evaluation of cationic polymer-coated 
nanocapsules as ocular drug carriers. Int f Pharm 153:41. 
84. De Campos A, Sanchez A and Alonso MJ (2001) Chitosan nanoparticles: A new vehicle 
for the improvement of the ocular retention of drugs. Application to cyclosporin A. Int 
J Pharm 224:159-168. 
85. De Campos A, Sanchez A and Alonso MJ (2003) The effect of a PEG vs a chitosan coating 
on the interaction of drug colloidal carriers with the ocular mucosa. Eur J Pharm Sci 
20:73-81. 
86. Calvo P, Vila-Jato JL and Alonso MJ (1997) Evaluation of cationic polymer-coated 
nanocapsules as ocular drug carriers. Int} Pharm 153:41-50. 
13 
Dendrimers as Nanoparticulate 
Drug Carriers 
SSnke Svenson and Donald A. Tomalia 
1. Introduction 
The development of molecular nanostructures with well-defined particle size and 
shape is of eminent interest in biomedical applications such as the delivery of active 
pharmaceuticals, imaging agents, or gene transfection. For example, constructs utilized 
as carriers in drug delivery generally should be in the nanometer range and 
uniform in size to enhance their ability to cross cell membranes and reduce the risk 
of undesired clearance from the body through the liver or spleen. Two traditional 
routes to produce particles that will meet some of these requirements have been 
widely investigated. The first route takes advantage of the ability of amphiphilic 
molecules (i.e. molecules consisting of a hydrophilic and hydrophobic moiety) to 
self-assemble in water above a system-specific critical micelle concentration (CMC) 
to form micelles. Size and shape of these micelles depend on the geometry of the 
constituent monomers, intermolecular interactions, and conditions of the bulk solution 
(i.e. concentration, ionic strength, pH, and temperature). Spherical micelles are 
monodisperse in size; however, they are highly dynamic in nature with monomer 
exchange rates in millisecond to microsecond time ranges. Micelles have the ability 
to encapsulate and carry lipophilic actives within their hydrocarbon cores. Depending 
on the specific system, some micelles either spontaneously rearrange to form 
liposomes after a minor change of solution conditions, or when they are exposed 
to external energy input such as agitation, sonication, or extrusion through a filter 
277 
278 Svenson &Tomalia 
membrane. Liposomes consist of bilayer lipid membranes (BLM) enclosing an aqueous 
core, which can be utilized to carry hydrophilic actives. Furthermore, liposomes 
with multilamellar membranes provide cargo space for lipophilic actives as well. 
However, most liposomes are considered energetically metastable, and will eventually 
rearrange to form planar bilayers.1'2 The second route relies on engineering the 
well-defined particles through processing protocols. Examples for this approach 
include (i) shearing or homogenization of oil-in-water (o/w) emulsions or w / o / w 
double emulsions to produce stable and monodisperse droplets, (ii) extrusion of 
polymer strands or viscous gels through nozzles of defined size to manufacture stable 
and monodisperse micro and nanospheres, (iii) layer-by-layer (LbL) deposition 
of polyelectrolytes and other polymeric molecules around colloidal cores, resulting 
in the formation of monodisperse nanocapsules after the removal of the templating 
core, and (iv) controlled precipitation from a solution into an anti-solvent, including 
supercritical fluids. Size, degree of monodispersity, and stability of these structures 
depend on the systems that are being used in these applications.3 These systems 
and their utilization in drug delivery are being discussed in detail in other chapters 
of this book. 
Currently, a new third route to create very well-defined, monodisperse, stable 
molecular level nanostructures is being studied based on the "dendritic state" 
architecture.4 Dendritic architecture is undoubtedly one of the most pervasive 
topologies observed throughout biological systems at virtually all dimensional 
length scales. This architecture is found at the meter scale in tree branching and 
roots, on the centimeter and millimeter scales in circulatory topologies in the human 
anatomy such as lungs, kidney, liver, and spleen, and on the micrometer scale in 
cerebral neurons. On the nanometer level, key examples of dendritic structures 
include glycogen, amylopectin, and proteoglycans. Amylopectins and glycogen 
are critical molecular level constructs involved in energy storage in plants and 
animals, while proteoglycans are an important constituent of connective tissue, 
determining its viscoelastic properties. Upon the analysis of these ubiquitous dendritic 
patterns, it is evident that these highly branched architectures offer unique 
interfacial and functional performance advantages. The objective of this review is to 
study the use of dendrimers in drug delivery applications. Four main properties of 
dendrimers will be discussed: (i) nanoscale container properties (i.e. encapsulation 
and transport of a drug), (ii) nano-scaffolding properties (i.e. surface adsorption 
or attachment of a drug and/or targeting ligand), (iii) dendrimers as drugs, and 
(iv) biocompatibility of dendrimers. In addition, routes of application currently 
investigated will be presented. Particular emphasis will be placed on poly 
(amidoamine) (PAMAM) dendrimers, the first and most extensively studied family 
of dendrimers.4c'5 
Dendrimers as Nanoparticulate Drug Carriers 279 
2. Nanoscale Containers — Micelles, Dendritic Boxes, 
Dendrophanes, and Dendroclefts 
Dendrimers may be visualized as consisting of three critical architectural domains: 
(i) the multivalent surface, containing a larger number of potentially reactive/ 
passive sites (nano-scaffolding), (ii) the interior shells (i.e. branch cell layers defined 
by dendrons) surrounding the core, and (iii) the core to which the dendrons are 
attached. The two latter domains represent well-defined nano-environments, which 
are protected from the outside by the dendrimer surface (nanoscale containers) 
in the case of higher generation dendrimers. These domains can be tailored for 
a specific purpose. The interior is well-suited for host-guest interaction and the 
encapsulation of guest molecules. 
2.1. Dendritic micelles 
Tomalia and coworkers demonstrated by electron microscopy observation that 
sodium carboxylated PAMAM dendrimers possess topologies reminiscent of regular 
classical micelles.4 It was also noted from electron micrographs that a large 
population of individual dendrimers possessed a hollow core. Supporting these 
observations, Turro and colleagues designed a hydrophobic 12-carbon atom alkylene 
chain into the core of a homologous series of PAMAM dendrimers (G = 2, 
3, and 4) to mimic the hydrophobic and hydrophilic core-shell topology of a regular 
micelle. The hosting properties of this series towards a hydrophobic dye as a 
guest molecule were then compared with a PAMAM dendrimer series possessing 
non-hydrophobic cores (e.g. NH3 and ethylenediamine). Dramatically enhanced 
emission of the hydrophobic dye was noted in aqueous solution in the presence 
of hydrophobic versus hydrophilic cored dendrimers.6a Less polar dendrimers 
(i.e. dendrimers containing aryl groups or other hydrophobic moieties as building 
elements), behave as inverse micelles.6b A critical property difference relative 
to micelles is the increased density of surface groups with higher generations. At 
some generational level, the surface groups will reach the so-called "de Gennes 
dense packing" limit and seal the interior from the bulk solution (Fig. I ) . 7 - 9 The 
limit depends on the strength of intramolecular interactions between adjacent surface 
groups, and therefore, on the condition of the bulk solution (i.e. pH, polarity 
and temperature). 
This nanoscale container feature, originally noted for PAMAM dendrimers 
by Tomalia et al. and referred to as "unimolecular encapsulation", can be utilized 
to tailor the encapsulation and release properties of dendrimers in drug delivery 
applications.910 For example, adding up to a limiting amount of Xmmol of 
either 2,4-dichlorophenoxyacetic acid or aspirin (acetylsalicyhc acid) to 1 mmol of 
280 Svenson & Tomalia 
,--oV..> jfc-iiw-'i.. J&:5Sgtfhv ;#88c?pi „•;*: 7.:;^- *#.# p f e $$8§? 
• • J • ^ • ' ' ^ ,s&$p?* '%$$? 
4 5 6 7 8 9 10 
Fig. 1. Periodic properties of PAMAM dendrimers generations G = 4-10, depicting the 
decreasing distances between surface charges (Z-Z). The "de Gennes dense packing" appears 
atG = 8. Dendrimers G = 4-6 display "nanoscale container" properties, the larger analogues 
G = 7-10 display "nano-scaffolding" properties. 
STARBURST® carbomethoxy-terminated PAMAM dendrimers generations 0.5-5.5 
produced spin-lattice relaxation times (Tj) much lower than the values of these 
guest molecules in solvent without dendrimer. The new relaxation times decreased 
for generations 0.5-3.5, but remained constant for generations 3.5 to 5.5. The maximum 
concentration X varied uniformly from 12 (generation 0.5) to 68 (generation 
5.5). On the basis of these maximum concentrations, the guest-to-host ratios were 
shown to be ~ 4:1 by weight and ~ 3:1 based on a molar comparison of dendrimer 
guest carboxylic acid-to-interior tertiary nitrogen moieties for generations 2.5-5.5. 
Exceeding the maximum concentration X resulted in the appearance of a second 
relaxation time, Tv, characteristic of the guest molecules in bulk solvent phase.10 
2.2. Dendritic box (Nano container) 
Surface-modification of G = 5 poly(propyleneimine) (PPI) dendrimers with 
Boc-protected amino acids induced dendrimer encapsulation properties by the 
formation of dense, hydrogen-bonded surface shells with solid-state character 
("dendritic box").8 Small guest-molecules were captured in such dendrimer interiors 
and were unable to escape even after extensive dialysis. The maximum amount 
of entrapped guest molecules was directly proportional to the shape and size of the 
guest molecules, as well as to the amount, shape and size of the available internal 
dendrimer cavities. Four large guest-molecules (i.e. Rose Bengal) and 8-10 small 
guest-molecules (i.e. p-nitrobenzoic acid) could be simultaneously encapsulated 
within PPI dendrimers containing four large and twelve smaller cavities. Remarkably, 
this dendritic box could be opened under controlled conditions to release either 
some or all of the entrapped guest molecules. For example, partial hydrolysis of 
the hydrogen-bonded Boc-shell liberated only small guest-molecules, whereas total 
hydrolysis released all sizes of entrapped molecules.8'11-12 
Although the "dendritic box" concept demonstrates the unique shapedependent 
cargo space that can be found in certain dendrimers, other parameters 
have to be considered as well for delivering and releasing therapeutic drugs 
Dendrimers as Nanoparticulate Drug Carriers 281 
under physiological conditions. From a thermodynamic perspective, free guestmolecules 
(i.e. drugs) can be distinguished from those encapsulated or bound in a 
complex by finite energy barriers related to the ease of entry and departure to the 
dendrimer cavities. If the drug molecule is incompatible with either the dimension 
or hydrophilic/lipophilic character of the dendrimer cavity, a complex might not 
form, or the guest might only be partially encapsulated within the dendrimer host. A 
hydrophobic drug would be expected to associate with a dendrimer core to achieve 
maximum contact with its hydrophobic domain. In addition, the hydrophobic character 
of this guest molecule would be expected to isolate itself from the dendrimer 
surface and the interface to the bulk solution to afford minimum contact with polar 
and aqueous domains (i.e. physiological media). Notably, the hydrophobic and 
hydrophilic properties, as well as other non-covalent binding properties of these 
spatial binding-sites are expected to strongly influence these guest-host relationships. 
Analysis of a typical symmetrically branched dendrimer makes it apparent 
that there are other subtle and yet important parameters that could control the interior 
space of a dendrimer and influence the guest-host interactions. These include 
components such as branching angles, branching symmetry rotational angles, and 
the length of a repeat-unit segment.13 Of equal importance are the properties of the 
core. Within a homologous PAMAM dendrimer series, the effect of changing the 
length scale of the core on dendrimer guest-host properties was studied. Specifically, 
a series of polyhydroxy-surfaced PAMAM dendrimers with core molecules 
differing in length by one carbon atom (NH2-Cn-NH2 with n = 2-6) were synthesized. 
Three aromatic carboxylic acids, differing systematically by one aromatic 
ring (benzoic acid, 1-naphthoic acid, 9-anthracene carboxylic acid), were examined 
as guest-molecule probes. Two sets of dendrimers, possessing 24 and 48 surface 
hydroxy groups, were investigated.14 The observed trends can be summarized as 
follows: (i) in general, all dendritic hosts accommodated larger amounts of the 
smaller guest-molecule (i.e. molar uptake benzoic > 1-naphthoic > 9-anthracene 
carboxylic acid). This observation was particularly significant for the more congested 
dendrimer surface having 48 surface OH-groups. (ii) Uptake maxima values 
specific to both the core size and the specific guest-probe were noted. This observation 
might be related to the combination of shape and lipophilicity manifested by the 
guest probe, (iii) A decrease in the molar uptake was measured for all probes as the 
core was enhanced beyond an ideal dimension (i.e. 5-6 carbons). It is therefore obvious 
that both core size and surface congestion dramatically affect the cargo-space 
of the dendrimer host. Furthermore, it is apparent that size and shape of the guest 
probe can significantly affect the maximum loading as a function of core size. Finally, 
it should also be noted that for the dendrimers G = 2 (24-OH) and G = 3 (48-OH), 
the guest probes had desirable release properties from the host as a function of time, 
when re-dissolved in water. Performing these same experiments using a dendrimer 
282 Svenson & Tomalia 
with more densely packed surface groups (i.e. G = 4 with 96 surface OH-groups) 
appeared to produce dendritic box behavior. Although guest molecules could be 
encapsulated within the core, the release from the host was delayed as determined 
by analysis after extensive dialysis.14 Structure-property relationships in dendritic 
encapsulation have been studied extensively, mainly using photoactive and redoxactive 
model dendrimers to gain a better understanding of the structural effects 
that cores and branches have on encapsulation.15-17 
2.3. Dendrophanes and dendroclefts 
Specific binding of guest molecules to the dendrimer core can affect the loading 
capacity by enhancing specific interactions between the core and guest (i.e. 
hydrophobic and polar interactions). Dendrimers specifically tailored to bind 
hydrophobic guests to the core have been created by Diederich and coworkers 
and coined "dendrophanes". These water-soluble dendrophanes are built around 
a cyclophane core, and can bind aromatic compounds, presumably via p -p interactions. 
Dendrophanes were shown to be excellent carriers of steroids.18'19 The same 
group synthesized dendrimers tailored to bind more polar bioactive compounds 
to the core, coined "dendroclefts".20'21 In another approach, the surface amines 
of PAMAM dendrimers were modified with tris(hydroxymethyl)aminomethane 
(TRIS) to create water-soluble dendrimers capable of binding carboxylic aromatic, 
antibacterial compounds, which could be released by lowering the pH.14 An alternative 
approach to creating dendritic hosts with highly selective guest recognition 
utilized the principle of "molecular imprinting".22 A dendrimer consisting of a porphyrin 
core and a surface containing terminal double bonds was polymerized into 
a polydendritic network. Subsequently, the base-labile ester bonds between cores 
and dendritic wedges were cleaved, releasing the porphyrin core from the dendritic 
polymer. This polymer was capable of selectively binding porphyrins with 
association constants of 1.4 x 105 M_1. Very recently, an impressive approach has 
been presented, using tandem mass spectrometry, i.e. the combination of electrospray 
ionization (ESI) and collision-induced dissociation (CID) mass spectrometers 
connected in series, to investigate the dynamic behavior of host-guest dendrimer 
complexes.23 This approach offers the potential to provide better insights into these 
constructs. 
3. Dendrimers in Drug Delivery 
Dendrimers have been utilized to carry a variety of small molecule pharmaceuticals 
with the purpose to enhance their solubility and therefore bioavailability, and to 
utilize the passive and active targeting properties of dendrimers, either through the 
Dendrimers as Nanoparticulate Drug Carriers 283 
"Enhanced Permeability and Retention" (EPR)24 effect or specific targeting ligands. 
Some aspects of dendrimers in drug delivery have been reviewed recently.13,25-27 
In the following, selected examples of important drug delivery aspects will be 
presented. 
3.1. Cisplatin 
Encapsulation of the well-known anticancer drug cisplatin within PAMAM dendrimers 
gives complexes that exhibit slower release, higher accumulation in solid 
tumors, and lower toxicity compared with free cisplatin.28'29 Cisplatin is an antitumor 
drug that exerts its effects by forming stable DNA-cisplatin complexes 
through intrastrand cross-links, resulting in an alteration of the DNA structure that 
prevents replication and activates cell repair mechanisms. The cell detects defective 
DNA and initiates apoptosis. Cisplatin is effective in treating several cancers such 
as ovarian, head and neck, and lung cancers, as well as melanomas, lymphomas, 
osteosarcomas, bladder, cervical, bronchogenic, and oropharyngeal carcinomas. 
Unfortunately, cisplatin has many adverse side effects to the body, the most important 
being nephrotoxicity and cytotoxicity to non-cancerous tissue, because of the 
non-selective interaction between cisplatin and DNA. In addition, the therapeutic 
effect of cisplatin is limited by its poor water solubility (1 mg/mL), low lipophilicity, 
and the development of resistance to cisplatin drugs. Although numerous cisplatin 
derivatives have undergone preclinical and clinical testing, only cisplatin and its 
derivatives carboplatin and oxaliplatin have been approved for routine clinical use 
(Fig. 2).30 
Preliminary studies gave cisplatin loadings of 15-25 wt% for PAMAM dendrimers 
generation 3.5 (size ~ 3.5 nm; MW ~ 13 kDa). In comparison, the cisplatin 
loading of linear poly(amidoamines) and linear N-(2-hydroxypropyl) methacrylamide 
(HPMA; MW 25-31 kDa) was found to be 5-10 wt% and 3-8 wt%, 
respectively. HPMA-cisplatin complexes are currently in clinical trials.31 The 
cisplatin-dendrimer complex could be visualized by Atomic Force Microscopy 
(AFM; carbon nanotip) as shown in Fig. 3. 
H3N, CI 
H3NT \ 
H3N- \ l ^ V 
O 
Fig. 2. Chemical structures of the platinum drugs cisplatin (PLATINOL®), carboplatin 
(PARAPLATIN®), and oxaliplatin (ELOXATIN™). 
284 Svenson & Tomalia 
Fig. 3. AFM images of cisplatin-dendrimer complexes at 120 (left) and 4nm (right) 
magnification. 
Table 1 AUC value (/xg Pt/mLblood or /xg Pt/organ) 
over 48 hours; 5 mice/data point. 
Organ Cisplatin Cisplatin-dendrimer Complex 
Tumor 5.3 25.4 
Blood 9.4 10.7 
Liver 51.6 17.0 
Kidney 57.6 138.1 
The tumor activity of the cisplatin-dendrimer formulation was studied using 
B16F10 cells. These cells were injected into C57 mice subcutaneously (s.c.) to provide 
a solid tumor model. After approximately 12 days, when the tumors had developed 
to a mean area of 50-100 mm2, the animals were injected i.v. with a single dose of 
either cisplatin or cisplatin-dendrimer complex (1 mg/kg cisplatin for both formulations). 
At certain time points within 48 hours, animals were culled and blood and 
tissue samples were taken. Compared with cisplatin alone, the cisplatin-dendrimer 
complex was found to accumulate preferentially in the tumor site relatively quickly 
after the injection. The tumor area under the curve (AUC) for the complex was 
5 times higher than that of free cisplatin, while that in the kidney only increased 
2.4 times, and accumulation in the liver was reduced (Table I).29 
Another recent study revealed a sufficient stability of cisplatin-dendrimer complexes, 
with a 20% release of cisplatin over the first 8 hours, and an additional 60% 
release within 150 hours. In vivo animal efficacy of the platinate was demonstrated 
using B16F10 tumor cells that are subcutaneous implanted into mice. The tumor 
was allowed to grow for 7 days prior to treatment with two doses of drug on day 
7 and day 14, providing equal cisplatin (5 mg/kg) doses in both the dendrimercisplatin 
complex and free cisplatin. A tumor weight reduction of ~ 40% above that 
observed for the free drug was found in this study. 
Dendrimers as Nanoparticulate Drug Carriers 285 
3.2. Silver salts 
The encapsulation of silver salts within PAMAM dendrimers produced conjugates, 
exhibiting slow silver release rates and antimicrobial activity against various Gram 
positive bacteria.32 PAMAM dendrimers, generation four with ethylenediamine 
(EDA) core and tris(2-hydroxymethyl)amidomethane (TRIS) OH-surface and generation 
five, EDA core with carboxylate COO~ surface, were used. Silver containing 
PAMAM complexes were prepared by adding aqueous solutions of the dendrimers 
to the calculated amount of silver acetate powder. Although CHaCOOAg is hardly 
soluble in water, it quickly dissolved in the PAMAM solutions. This enhancement 
is due to the combined action of the silver carboxylate salt formation and/or to the 
complex formation with the internal dendrimer nitrogens. This procedure resulted 
in slightly yellow dendrimer-complex/salt solutions that very slowly photolyzed 
when exposed to light, into dark brown, metallic silver, containing dendrimersilver 
nanocomposite solutions. Final sample concentrations were confirmed by 
atomic absorption spectroscopy. For antimicrobial testing, the standard agar overlay 
method was used. In this test, dendrimer-silver compounds were examined 
for diffusible antimicrobial activity by placing a lO-^L sample of each solution 
onto a 6-mm filter paper disk and applying the disk to a dilute population of the 
test organisms, Staphylococcus aureus, Pseudomonas aeruginosa, and Escherichia coli. 
The silver-dendrimer complexes displayed antimicrobial activity, comparable to or 
better than those of silver nitrate solutions. Interestingly, increased antimicrobial 
activity was observed with dendrimer carboxylate salts, which was attributed to 
the very high local concentration (256 carboxylate groups around a 5.4 nm diameter 
sphere) of nanoscopic size silver composite particles that are accessible for 
microorganisms. The antimicrobial activity was smaller when internal silver complexes 
were applied instead of silver adducts to the surface, indicating that the 
accessibility of the silver is an important factor. 
3.3. Adriamycin, methotrexate, and 5-fluorouracil 
The anticancer drugs, adriamycin and methotrexate, were encapsulated into generations 
3 and 4 PAMAM dendrimers which had poly(ethylene glycol) monomethyl 
ether chains with molecular weights of 550 and 2000 Da attached to their surfaces 
via urethane bonds (Fig. 4). The encapsulation efficiency was dependent on the 
PEG chain length and the size of the dendrimer, with the highest encapsulation 
efficiencies (on average, 6.5 adriamycin molecules and 26 methotrexate molecules 
per dendrimer) found for the G = 4 PAMAM terminated with PEG2000 chains. 
The drug release from this dendrimer was sustained at low ionic strength, again 
reflecting PEG chain length and dendrimer size, but fast in isotonic solution.33 In a 
related study, it was reported that the surface coverage of PAMAM dendrimers with 
286 Svenson & Tomalia 
Fig. 4. Above: Structures of anticancer drugs adriamycin (left) and methotrexate (right). 
Below: Schematic presentations of the encapsulation of methotrexate (left) and 5-fluorouracil 
(right) into PAMAM dendrimers. 
PEG2000 chains had little influence on the encapsulation efficiency of methotrexate, 
but affected the release rate.34 
A similar construct between PEG chains and PAMAM was utilized to deliver the 
anticancer drug 5-fluorouracil. Encapsulation of 5-fluorouracil into G — 4 PAMAM 
dendrimers with carboxymethyl PEG5000 surface chains revealed reasonable drug 
loading, a reduced release rate, and reduced hemolytic toxicity compared to the 
non-PEGylated dendrimer (Fig. 4).35 
3.4. Etoposide, mefenamic acid, diclofenac, and venlafaxine 
The combination between dendrimers and hydrophilic and/or hydrophobic polymer 
chains has recently been extended to solubilize the hydrophobic anticancer 
drug etoposide. A star polymer composed of amphiphilic block copolymer arms 
has been synthesized and characterized. The core of the star polymer was a 
generation two PAMAM-OH dendrimer, the inner block of the arm a lipophilic 
poly(e-caprolactone) (PCL) and the outer block of the arm a hydrophilic PEG500o- 
The star-PCL polymer was synthesized first by ring-opening polymerization of 
e-caprolactone with the PAMAM-OH dendrimer as initiator. The PEG polymer 
was then attached to the PCL terminus by an ester-forming reaction. Characterization 
with SEC, 1-H NMR, FTIR, TGA, and DSC confirmed the star structure of the 
polymers. A loading capacity of up to 22% (w/w) was achieved with etoposide. 
Dendrimers as Nanoparticulate Drug Carriers 287 
A cytotoxicity assay demonstrated that the star-PCL-PEG copolymer was nontoxic 
in cell culture.36 
Citric acid-poly(ethylene glycol)-citric acid (CPEGC) triblock dendrimers generations 
1-3 were applied to encapsulate small molecule drugs such as mefenamic 
acid and diclofenac. The formulations were stored at room temperature for up to 
ten months and remained stable with no reported release of the drugs.37 
The attachment of the novel third-generation antidepressant venlafaxine onto 
anionic PAMAM dendrimers (G = 2.5) via a hydrolyzable ester bond and the incorporation 
of this drug-dendrimer complex into a semi-interpenetrating network of 
an acrylamide hydrogel has been studied as a novel drug delivery formulation to 
avoid the currently necessary multiple daily administration of the antidepressant. 
The effect of PEG concentration and molecular weight was studied to find optimal 
release conditions.38 
3.5. Ibuprofen, indomethacin, nifedipine, naproxen, paclitaxel, 
and methylprednisolone 
The anti-inflammatory drug ibuprofen was used as a model compound to study 
its complexation and encapsulation into generations 3 and 4 PAMAM dendrimers 
and a hyperbranched polyester, having approximately 128 surface OH-groups. It 
was found that up to 78 ibuprofen molecules were complexed by the PAMAM dendrimers 
through electrostatic interactions between the dendrimer amines and the 
carboxyl group of the drug. In contrast, up to 24 drug molecules were encapsulated 
into the hyperbranched polyol.39 The drug was successfully transported into A549 
human lung epithelial carcinoma cells by the dendrimers. The PAMAM dendrimers 
with either amino or hydroxy surfaces entered the cells faster (in approximately 1 hr) 
than the hyperbranched polyol (approximately 2 hrs). However, both entries were 
faster than the pure drug. The anti-inflammatory effect of ibuprofen-dendrimer 
complexes was demonstrated by more rapid suppression of COX-2 mRNA levels 
than that achieved by the pure drug.40 
The non-steroidal anti-inflammatory drug (NSAID) indomethacin is practically 
insoluble in water and only sparingly soluble in alcohol. Encapsulation of 
indomethacin into generation 4 PAMAM dendrimers with amino, hydroxy, and 
carboxylate surfaces remarkably enhanced the drug solubility in water, and therefore, 
its bioavailability (Fig. 5).41 The encapsulation efficiency of indomethacin into 
PAMAM dendrimers is dependent on the dendrimer size (G6 > G5 > G4 > G3) 
and the surface functionalization, (NH2 > PEG = PYR > AE) (Fig. 6).42 
The effect of PAMAM dendrimer generation size and surface functional group 
on the aqueous solubility, and therefore, bioavailability of the calcium channel 
blocking agent nifedipine has been studied using PAMAM dendrimers with EDA 
288 Svenson & Tomalia 
CH, 
COOH 
E 800^ 
600- 
g 400- 
200 
0.1 0.2 
Dendrimerconc. (%v/w) 
Fig. 5. Molecular structure of indomethacin and its solubility profiles in the presence of differing 
concentrations of G4-NH2/ (•) G4-OH (•), and G4.5-COOH (A) PAM AM dcndrimers 
at pH 7 (« = 3, R.S.D. < 5%). 
Encapsulation efficiency of EDA core PAMAM dendrimer 
« 
E 
80 - 
60- 
g 40 
3 20- 
I 
ft 
,rf sJl 
• G3 
• G4 
• G5 
• G6 
—n .r-dl ^ 
NH2 PEG PYR AE COONa 
Surface Functionality 
sue TRIS 
Fig. 6. Encapsulation efficiency into PAMAM dendrimers generations 3-6 with amino 
(NH2), poly(ethylene glycol) (PEG), carbomethoxypyrrolidinone (PYR), amidoethanol 
(AE), sodium carboxylate (COONa), succinamic acid (SUC), and tris(hydroxymethyl)- 
aminomethane (TRIS) surface groups. 
core and amino surface (G = 0,1,2,3) or ester surface (G — 0.5,1.5,2.5) at pH 4, 
7 and 10. The solubility enhancement of nifedipine was higher in the presence of 
ester-terminated dendrimers than their amino-terminated analogues, possessing 
the same number of surface groups. The nifedipine solubility expectedly increased 
with the size of the dendrimers. For pH 7, the sequence G2.5 > G3 > G1.5 > G2 > 
G0.5 > Gl > GO was reported.43 
In another approach, the non-steroidal anti-inflammatory drug naproxen was 
covalently attached to unsymmetrical poly(arylester) dendrimers to prepare a complex 
with enhanced water solubility of the drug and access for hydrolytic cleavage 
Dendrimers as Nanoparticulate Drug Carriers 289 
1E-3 0.01 0.1 1 0 2k 4k 6k 8k 
Concentration (M) Molecular weight 
Fig. 7. Aqueous paclitaxel solubility as a function of the polyglycerol dendrimer concentration 
(mean ± SD, n = 3); G5 (circle), G4 (triangle), G3 (square), and PEG400 (diamond) 
(left). Molecular weight dependency of dendrimers (closed circle) and PEG (open circle) on 
the aqueous paclitaxel solubility. The concentration of dendrimers and PEG was 10wt%. 
(Reproduced with permission from Ref. 45. Copyright 2004 American Chemical Society.) 
of the bond between drug and carrier. Detailed results on the biological evaluation 
of these complexes have not been reported.44 
The anticancer drug paclitaxel, which is being used to treat metastatic breast 
and ovarian cancers and Kaposi's sarcoma, has poor water solubility. To enhance 
its bioavailability, paclitaxel has been encapsulated into polyglycerol dendrimers, 
resulting in a 10,000-fold improved water solubility compared with the pure drug, 
which is much higher than that found for PEG400, a commonly used linear chain 
cosolvent or hydrotropic agent (Fig. 7). The drug release rate was a function of the 
dendrimer generation.45 
Generation 4 PAMAM dendrimers with hydroxy surface have been utilized 
to improve the bioavailability of the corticosteroid methylprednisolone, which 
decreases inflammation by stabilizing leukocyte lysosomal membrane. By connecting 
the drug to the dendrimer using glutaric acid as the spacer, a payload of 32 
wt% was achieved. The drug-dendrimer complex was taken up by A549 human 
lung epithelial carcinoma cells and mostly localized in the cytosol. The complex 
showed a pharmacological activity comparable to the free drug as measured by the 
inhibition of the prostaglandin secretion.46 
3.6. Doxorubicin and camptothecin — self-immolative dendritic 
prodrugs 
An exciting new approach to dendritic drug delivery involves the utilization of a 
drug as a part of the dendritic molecule. Self-immolative dendrimers have recently 
been developed and introduced as a potential platform for a multi-prodrug. These 
unique structural dendrimers can release all of their outer branch units through 
10 
•jj 0.01- 
I 1E-3, 
290 Svenson & Tomalia 
co3 
>~q 
vr 
HsO 
CegS^-'y 
«Ht 
Fig. 8. Mechanism of dimeric prodrug activation by a single enzymatic cleavage. (Reproduced 
with permission from Ref. 47. Copyright 2004 American Chemical Society.) 
a self-immolative chain fragmentation, initiated by a single cleavage at the dendrimer's 
core. Incorporation of drug molecules as these outer branch units and an 
enzyme substrate as the trigger can generate a multi-prodrug unit that will be activated 
with a single enzymatic cleavage (Fig. 8). The first generation of dendritic 
prodrugs with doxorubicin and camptothecin as branch units and retro-Michael 
focal trigger, which can be cleaved by the catalytic antibody 38C2, has been reported. 
Bioactivation of the dendritic prodrugs was evaluated in cell-growth inhibition 
assay with the Molt-3 leukemia cell line in the presence and absence of antibody 
38C2. A remarkable increase in toxicity was observed. Dependent on the linker 
molecule, different numbers of drug molecules can be released in one single activation 
step.47'48 
In a more "classical" approach to deliver doxorubicin, two polyester-based 
dendrimers (generation 4 with trisphenolic core) were synthesized, one carrying a 
hydroxy surface, the other a tri(ethylene glycol) monomethyl ether surface. These 
dendrimers were compared with a 3-arm poly(ethylene oxide) star polymer, carrying 
G = 2 dendritic polyester units at the surface. The star polymer gave the most 
promising results regarding cytotoxicity and systemic circulatory half-life (72hrs). 
Therefore, the anticancer drug doxorubicin was covalently bound to this carrier via 
an acid-labile hydrazone linkage. The cytotoxicity of doxorubicin was significantly 
reduced (80-98%) and the drug was successfully taken up by several cancer cell 
lines.49 
Dendrimers as Nanoparticulate Drug Carriers 291 
3.7. Photodynamic therapy (PDT) and boron neutron capture 
therapy (BNCT) 
Dendrimers have been used to optimize the antitumor effect in photodynamic 
therapy (PDT) and boron neutron capture therapy (BNCT). One of the newest 
developments in the dendrimer field is their application to photodynamic therapy 
(PDT). This cancer treatment involves the administration of a light-activated 
photosensitizing moiety that selectively concentrates in diseased tissue. Subsequent 
activation of the photosensitizer leads to the generation of reactive oxygen, primarily 
singlet oxygen, that damages intracellular components such as lipids and amino 
acid residues through oxidation, ultimately leading to cell death by apoptosis. 
Disadvantages of currently used photosensitizers include skin phototoxicity, poor 
selectivity for tumor tissue, poor water solubility, and difficulties in the treatment 
of solid tumors because of the impermeability of the skin and tissues to the visible 
light required to excite the chromophores. 
In one set of studies, dendrimers have been constructed around a light harvesting 
core (i.e. a porphyrin).50 To reduce the toxicity under non-irradiative conditions 
(dark toxicity) and to prevent aggregation, and consequently, self-quenching of the 
porphyrin cores, these dendrimers have been further encapsulated into micelles. 
For example, poly(ethylene glycol)-b-poly(aspartic acid) and PEG-b-poly(L-lysine) 
micelles have been studied in this regard. These micelles are stable under physiological 
conditions pH 6.2 to 7.4. However, they disintegrate in the acidic intracellular 
endosomal compartment (pH ~ 5.0).51/52 Alternatively, the photosensitizer 
5-aminolevulinic acid has been attached to the surface of dendrimers and studied 
as an agent for PDT of tumorigenic keratinocytes.53 Photosensitive dyes have been 
incorporated into dendrimers and utilized in PDT devices. For example, uptake, 
toxicity, and the mechanism of photosensitization of the dye pheophorbide a (pheo) 
was compared with its complex with diaminobutane poly(propylene imine) (DAB) 
dendrimers in human leukemia cells in vitro.5i 
The second therapy, boron neutron capture therapy, is a cancer treatment based 
on a nuclear capture reaction. When 10B is irradiated with low energy or thermal 
neutrons, highly energetic a-particles and 7Li ions are produced, that are toxic to 
tumor cells. To achieve the desired effects, it is necessary to deliver 10B to tumor cells 
at a concentration of at least 109 atoms per cell. High levels of boron accumulation 
in tumor tissue can be achieved by using boronated antibodies that are targeted 
towards tumor antigens. However, this approach can impair the solubility and 
targeting efficiency of the antibodies. 
One study, involving intratumoral injection of a conjugation between a generation 
5 PAMAM dendrimer carrying 1100 boron atoms and cetuximab, a monoclonal 
antibody specific for the EGF receptor, showed that the conjugate was present 
292 Svenson & Tomalia 
Fig. 9. Schematic presentation of an EDA core G = 3 PAMAM dendrimer (1), the boron 
carrier Na(CH3)3NB10H8NCO (2), and the targeting ligand folic acid (3). (Reproduced with 
permission from Ref. 56. Copyright 2003 American Chemical Society.) 
at an almost 10-fold higher concentration in brain tumors than in normal brain 
tissue.55 To reduce the liver uptake observed for boronated PAMAM dendrimer 
conjugates, PEG chains were attached onto the dendrimer surface, in addition to 
the borane clusters, to provide steric shielding. As compared with a dendrimer without 
PEG chains, the amount of liver uptake was found to be less for PEG-conjugated 
dendrimers with an average of 1.0-1.5 chains of PEG2000/ but higher for dendrimers 
with 11 chains of PEG550. Folic acid moieties were also conjugated to the ends of 
the PEG chains to enhance the uptake of the dendrimers by tumors overexpressing 
folate receptors. Although this strategy was successful in enhancing localization 
of the molecules to tumors in mice bearing 24JK-FBP tumors expressing the folate 
receptor, it also led to an increase in the uptake of the dendrimers by the liver and 
kidneys.56 
4. Nano-Scaffolds for Targeting Ligands 
The surface of dendrimers provides an excellent platform for the attachment of cellspecific 
ligands, solubility modifiers, stealth molecules, reducing the interaction 
with macromolecules from the body defense system, and imaging tags. The ability 
to attach any or all of these molecules in a well-defined and controllable manner 
onto a robust dendritic surface, clearly differentiates dendrimers from other carriers 
such as micelles, liposomes, emulsion droplets, and engineered particles. 
4.1. Folic acid 
One example of cell-specific dendritic carriers is a dendrimer modified with folic 
acid. The membrane-associated high affinity folate receptor (hFR) is a folate binding 
protein that is overexpressed on the surface of a variety of cancer cells, and 
Dendrimers as Nanoparticulate Drug Carriers 293 
therefore, folate-modified dendrimers would be expected to internalize into these 
cells preferentially over normal cells via receptor-mediated endocytosis. Folatedendrimer 
conjugates have been shown to be well-suited for targeted, cancerspecific 
drug delivery of cytotoxic substances.56-59 
In a very recent study, branched poly(L-glutamic acid) chains were centered 
around PAMAM dendrimers generations 2 and 3 and poly(ethylene imine) (PEI) 
cores to create new biodegradable polymers with improved biodistribution and targeting 
ability. These constructs were surface-terminated with poly(ethylene glycol) 
chains to enhance their biocompatibility, and folic acid ligands to introduce cellspecific 
targeting. Cell binding studies have been performed using the epidermal 
carcinoma cell line, KB.60 
4.2. Carbohydrates 
In addition to folates, carbohydrates constitute another important class of biological 
recognition molecules, displaying a wide variety of spatial structures due to 
their branching possibility and anomericity. To achieve sufficiently high binding 
affinities between simple mono- and oligosaccharide ligands and cell membrane 
receptors, these ligands have to be presented to the receptors in a multivalent or 
cluster fashion.61,62 The highly functionalized surface of dendrimers provides an 
excellent platform for such presentations. The design, synthesis, and biomedical 
use of glycodendrimers, as well as their application in diagnostic and for vaccinations, 
have been thoroughly reviewed recently.63-69 For example, the Thomsen- 
Friedenreich carbohydrate antigen (T-antigen), /J-Gal-(l-3)-a-GalNAc, which has 
been well documented as an important antigen for the detection and immunotherapy 
of carcinomas, especially relevant to breast cancer, has been attached to the 
surface of PAMAM and other dendrimers.70-72 An enhanced binding affinity was 
observed for all glycodendrimers. These constructs could have potential in blocking 
the metastatic sites of invasive tumor cells. A series of dendritic ,6-cyclodextrin 
derivatives, bearing multivalent mannosyl ligands, has been prepared and their 
binding efficiency towards the plant lectin concanavalin A (Con A) and a mammalian 
mannose-specific cell surface receptor from macrophages has been studied. 
The effects of glycodendritic architecture on binding efficiency, molecular inclusion, 
lectin-binding properties, and the consequence of complex formation using 
the anticancer drug docetaxel on biological recognition were investigated.73 Di- to 
tetravalent dendritic galabiosides, carrying (Galal-4Gal) moieties on their surfaces, 
were studied as inhibitors of pathogens based on bacterial species such as E. colt and 
Streptococcus suis. Attachment of dendritic galabiosides onto cell surfaces would be 
expected to inhibit the attachment of bacteria using the same sugar ligand-receptor 
interactions. The study revealed a clear enhancement of the binding affinity between 
294 Svenson & Tomalia 
glycodendrons and cell surfaces, with an increasing number of sugar moieties.74 In 
a similar approach, glycodendrons carrying two to four /i-D-galactose moieties on 
their surface, while the dendron core was connected to a protein-degrading enzyme, 
were synthesized. These glycodendriproteins are expected to attach to the surface of 
bacteria, allowing the enzyme to degrade the bacterial adhesin, hence rendering the 
bacteria incapable of attaching to the cell surfaces.75 Anionic PAMAM dendrimers 
(G = 3.5) were conjugated to D(+)-glucosamine and D(+)-glucoseamine 6-sulfate. 
These water-soluble conjugates not only revealed immuno-modulatory and antiangiogenic 
properties, but synergistically prevented scar tissue formation after glaucoma 
filtration surgery. In a validated and clinically relevant rabbit study, the longterm 
success rate was increased from 30 to 80% using these dendrimer-conjugates.76 
4.3. Antibodies and biotin-avidin binding 
Generation 5 PAMAM dendrimerendrimers with amino surface were conjugated 
to fluorescein isothiocyanate as a means to analyze cell binding and internalization. 
Two different antibodies, 60bca and J591, which bind to CD14 and prostate-specific 
membrane antigen (PSMA) respectively, were used as model targeting molecules. 
The binding of the antibody-conjugated dendrimers to antigen-expressing cells 
was evaluated by flow cytometry and confocal microscopy. The conjugates specifically 
bound to the antigen-expressing cells in a time- and dose-dependent fashion, 
with affinity similar to that of the free antibody (Fig. 10). Confocal microscopic 
analysis suggested at least some cellular internalization of the dendrimer conjugate. 
Dendrimer-antibody conjugates are, therefore, a suitable platform for targeted 
molecule delivery into antigen-expressing cells.77 
Monolayers formed by generation 4 PAMAM dendrimers on a gold surface 
were functionalized with biotin and produced a biomolecular interface that was 
Fig. 10. Confocal microscopic analysis of HL60 cells, which were incubated (1 h at 4°C) 
with 12.5nM G5 PAMAM carrying fluorescence dye and 60bca antibody on the surface. The 
cells were rinsed and confocal images were taken. The left and right panels represent the FITC 
fluorescence and light images taken in the same cell. The arrow indicates the binding of the 
conjugate on the cell surface at 4°C. (Reproduced with permission from Ref. 77. Copyright 
2004 American Chemical Society.) 
Dendrimers as Nanoparticulate Drug Carriers 295 
capable of binding high levels of avidin. Avidin binding as high as 88% coverage 
of the surface was observed despite conditions that should cause serious steric 
hindrance. These dendritic monolayers were utilized as a model to study proteinligand 
interactions.78 
4.4. Penicillins 
The surfaces of PAMAM dendrimers, generations 0 to 3, were decorated with benzylpenicillin 
in an attempt to develop a new in vitro test to quantify IgE antibodies 
to specific ^-lactam conjugates, with the goal of improving the existing methods for 
diagnosing allergy to this type of antibiotic. The monodispersity of dendrimers is 
advantageous over conventional peptide carrier conjugates such as human serum 
albumin (non-precise density of haptens in their structure) and poly-L-lysine (mixture 
of heterogeneous molecular weight peptides). Preliminary radioallergosorbent 
tests (RAST), using sera from patients allergic to penicillin, have confirmed the usefulness 
of penicilloylated dendrimers.79 
Penicillin V was used as a model drug containing a carboxylic group and 
attached to the surface of PAMAM dendrimers generations 2.5 and 3, both containing 
32 surface functionalities. The drug was complexed to the dendrimers via 
amide or ester bonds. It was found in tests using a single-strain bacterium, Staphylococcus 
aureus, that the bioavailability of the penicillin was unaltered after the drug 
was released from the complex through ester bond hydrolysis.80 
5. Dendrimers as Nano-Drugs 
Dendrimers have been studied as antitumor, antiviral and antibacterial drugs.25 
The most prominent and advanced example is the use of poly(lysine) dendrimers, 
modified with sulfonated naphthyl groups, as antiviral drugs against the herpes 
simplex virus.81 Such a conjugate based on dendritic poly(lysine) scaffolding is 
VivaGel™, a topical agent currently under development by Starpharma Ltd., Melbourne, 
Australia, that can potentially prevent/reduce transmission of HIV and 
other sexually transmitted diseases (STDs). VivaGel™ (SPL 7013) is being offered 
as a water-based gel, with the purpose to prevent HIV from binding to cells in the 
body. The gel differs from physical barriers to STDs such as condoms, by exhibiting 
inhibitory activity against HIV and other STDs. In July 2003, following submission 
of an Investigational New Drug (IND) application, Starpharma gained clearance 
under U.S. FDA regulations to proceed with a Phase I clinical study to assess 
the safety of VivaGel™ in healthy human subjects. This phase 1 study, representing 
for the first time a dendrimer pharmaceutical tested in humans, compared 36 
women who received either various intra-vaginal doses of VivaGel™ or a placebo 
gel daily for one week. The trial was double blinded so that the volunteers, principal 
296 Svenson & Tomalia 
investigator and Starpharma did not know who was receiving placebo or VivaGel™. 
Study participants were assessed for possible irritant effects of the gel. Additionally, 
the women were assessed for any possible effect upon vaginal microflora (natural 
micro-organisms in the vagina) or absorption into the blood of the active ingredient 
of VivaGel™. A thorough review of the complete data revealed no evidence of irritation 
or inflammation. Preclinical development studies had demonstrated that 
VivaGel™ was 100% effective at preventing infection of primates exposed to a 
humanized strain of simian immunodeficiency virus (SHIV).82 In earlier studies, 
it was found that PAMAM dendrimers covalently modified with naphthyl sulfonate 
residues on the surface, also exhibited antiviral activity against HIV. This 
dendrimer-based nano-drug inhibited early stage virus/cell adsorption and later 
stage viral replication, by interfering with reverse transcriptase and/or integrase 
enzyme activities.83,84 
The general mode of action of antibacterial dendrimers is to adhere to and 
damage the anionic bacterial membrane, causing bacterial lysis.25,85 PPI dendrimers 
with tertiary alkyl ammonium groups attached to the surface have been shown to 
be potent antibacterial biocides against Gram positive and Gram negative bacteria. 
The nature of the counterion is important, as tetraalkylammonium bromides 
were found to be more potent antibacterials over the corresponding chlorides.86 
Poly(lysine) dendrimers with mannosyl surface groups are effective inhibitors of 
the adhesion of E. coli to horse blood cells in a haemagglutination assay, making 
these structures promising antibacterial agents.87 Chitosan-dendrimer hybrids 
have been found to be useful as antibacterial agents, carriers in drug delivery systems, 
and in other biomedical applications. Their behavior have been reviewed 
very recently88 Triazine-based antibiotics were loaded into dendrimer beads at 
high yields. The release of the antibiotic compounds from a single bead was sufficient 
to give a clear inhibition effect.89 In many cases, dendritic constructs were 
more potent than analogous systems based on hyperbranched polymers. 
The anti-prion activity of cationic phosphorus-containing dendrimers with tertiary 
amine surface groups has been evaluated. These molecules had a strong anti 
prion activity at non-toxic doses. They have been found to decrease the amount of 
pre-existing PrPSc from several prion starins, including the BSE strain. In addition, 
these dendrimers were able to reduce PrPSc accumulation in the spleen by more 
than 80%.90 
6. Routes of Application 
Most commonly, dendrimers are applied as parenteral injections, either directly 
into the tumor tissue or intravenous for systemic delivery. However, recent oral 
drug delivery studies using the human colon adenocarcinoma cell line, Caco-2, 
Dendrimers as Nanoparticulate Drug Carriers 297 
have indicated that low generation PAMAM dendrimers cross cell membranes, 
presumably through a combination of two processes, i.e. paracellular transport 
and adsorptive endocytosis.91 The P-glycoprotein (P-gp) efflux transporter does 
not effect dendrimers, and therefore, drug-dendrimer complexes are able to bypass 
the efflux transporter.92 
Furthermore, recent work has shown that PAMAM dendrimers enhanced the 
bioavailability of indomethacin in transdermal delivery applications.93 Similarly, 
the drug tamsulosin was used as a model to study transdermal delivery utilizing 
PAMAM dendrimers. The dendrimers were found to be weak penetration 
enhancers.94 However, no dendrimer-driven effect was observed for the drugs ketoprofen 
and clonidine. As an explanation, dendrimer-triggered drug crystallization 
within the transdermal delivery matrix was discussed, allowing the formation of 
drug polymorphs that can or cannot facilitate transdermal delivery95 
Several PAMAM dendrimers (generations 1.5, 2-3.5 and 4) with amine, carboxylate 
and hydroxyl surface groups were studied for controlled ocular drug 
delivery. The duration of residence time was evaluated after solubilization of these 
dendrimers in buffered phosphate solutions containing 2 parts per thousand (w/v) 
of fluorescein. The New Zealand albino rabbit was used as an in vivo model for qualitative 
and quantitative assessment of ocular tolerance and retention time, after a single 
application of 25 /xL of dendrimer solution to the eye. The same model was also 
used to determine the prolonged miotic or mydriatic activities of dendrimer solutions, 
some containing pilocarpine nitrate and some tropicamide, respectively. Residence 
time was longer for the solutions containing dendrimers with carboxylic and 
hydroxyl surface groups. No prolongation of remanence time was observed when 
dendrimer concentration (0.25-2%) increased. The remanence time of PAMAM dendrimer 
solutions on the cornea showed size and molecular weight dependency. This 
study allowed novel macromolecular carriers to be designed with prolonged drug 
residence time for the ophthalmic route.96 
7. Biocompatibility of Dendrimers 
Dendrimers have to exhibit low toxicity and be non-immunogenic in order to be 
widely used in biomedical applications. To date, the cytotoxicity of dendrimers 
has been primarily studied in vitro, however, a few in vivo studies have been 
published.25 As observed for other cationic macromolecules, including liposomes 
and micelles, dendrimers with positively charged surface groups are prone to destabilize 
cell membranes and cause cell lysis. For example, in vitro cytotoxicity IC50 
measurements (i.e. the concentration where 50% of cell lysis is observed) for aminoterminated 
PAMAM dendrimers revealed significant cytotoxicity on human intestinal 
adenocarcinoma Caco-2 cells.97,98 Furthermore, the cytotoxicity was found to be 
298 Svenson & Tomalia 
generation-dependent, with higher generation dendrimers being the most toxic. ' 
A similar generation dependence of amino-terminated PAMAM dendrimers was 
observed for the haemolytic effect, studied on a solution of rat blood cells.100 However, 
some recent studies have shown that amino-terminated PAMAM dendrimers 
exhibit lower toxicity than more flexible amino-functionalized linear polymers perhaps 
due to lower adherence of the rigid globular dendrimers to cellular surfaces. 
The degree of substitution, as well as the type of amine functionality, is important, 
with primary amines being more toxic than secondary or tertiary amines." 
Amino-terminated PPI and PAMAM dendrimers behave similarly with regard to 
cytotoxicity and haemolytic effects, including the generation-dependent increase 
of both.100'101 
Comparative toxicity studies on anionic (carboxylate-terminated) and cationic 
(amino-terminated) PAMAM dendrimers using Caco-2 cells have shown a significantly 
lower cytotoxicity of the anionic compounds.97 In fact, lower generation 
PAMAM dendrimers possessing carboxylate surface groups show neither haematotoxicity 
nor cytotoxicity at concentrations up to 2 mg/ml.100 The biocompatability 
of dendrimers is not solely determined by the surface groups. Dendrimers containing 
an aromatic polyether core and anionic carboxylate surface groups have shown 
to be haemolytic on a solution of rat blood cells after 24hrs. It is suggested that 
the aromatic interior of the dendrimer may cause haemolysis through hydrophobic 
membrane contact.100 
One way to reduce the cytotoxicity of cationic dendrimers may reside in partial 
surface derivatization with chemically inert functionalities such as PEG or fatty 
acids. The cytotoxicity towards Caco-2 cells can be reduced significantly (from 
IC50 ~ 0.13 mM to >lmM) after such a modification. This observation can be 
explained by the reduced overall positive charge of these surface-modified dendrimers. 
Apartial derivatization with as few as six lipid chains or four PEG chains on 
a G4-PAMAM, respectively, was sufficient to lower the cytotoxicity substantially.98 
In studies conducted at Dendritic Nano Technologies, Inc. using Caco-2 and two 
other cell lines, it was found that besides (partial) PEGylation of the surface, surface 
modification with pyrrolidone, another biocompatible compound, can significantly 
reduce cytotoxicity to levels far better than those of currently available products.102 
In some cases, the cytotoxicity of PAMAM dendrimers could be reduced by additives 
such as fetal calf serum.103 
Only a few systematic studies on the in vivo toxicity of dendrimers have been 
reported so far. Upon injection into mice, doses of 10 mg/kg of PAMAM dendrimers 
(up to G = 5), displaying either unmodified or modified amino-terminated surfaces, 
did not appear to be toxic.81-104 Hydroxy- or methoxy-terminated dendrimers 
based on a polyester dendrimer scaffold have been shown to be of low toxicity 
both in vitro and in vivo. At very high concentrations (40 mg/ml), these polyester 
Dendrimers as Nanoparticulate Drug Carriers 299 
dendrimers induced some inhibition of cell growth in vitro, but no increase in cell 
death was observed. Upon injection into mice, no acute or long-term toxicity problems 
were observed. The non-toxic properties make these new dendritic motifs very 
promising candidates for drug delivery devices.49 
Initial immunogenicity studies performed on unmodified amino-terminated 
PAMAM dendrimers showed no or weak immunogenicity of the G3-G7 dendrimers. 
However, later studies indicated some immunogenicity of these dendrimers, 
which could be reduced by surface-modification utilizing PEG chains.105 
8. Conclusions 
The high level of control over the architecture of dendrimers, their size, shape, 
branching length and density, and their surface functionality, makes these compounds 
ideal carriers in drug delivery applications. The bioactive agents may either 
be encapsulated into the interior of the dendrimers or they may be chemically 
attached or physically adsorbed onto the dendrimer surface, with the option to 
tailor the properties of the carrier to the specific needs of the active material and its 
therapeutic applications. Furthermore, the high density of surface groups allows 
attachment of targeting groups as well as groups that modify the solution behavior 
or toxicity of dendrimers. Surface-modified dendrimers themselves may act 
as nano-drugs against tumors, bacteria and viruses. This review of drug delivery 
applications of dendrimers clearly illustrates the potential of this new "fourth architectural 
class of polymers"106 and substantiates the high optimism for the future of 
dendrimers in this important field. 
Acknowledgments 
The authors wish to thank all contributors to this fascinating field of research, as 
well as the funding agents that have supported this work over the years. In particular, 
DNT would like to acknowledge current funding by the US Army Research 
Laboratory (ARL) (Contract # W911NF-04-2-0030). 
References 
1. Svenson S (2004) Controlling surfactant self-assembly. Curr Opin Coll Interj Sci 9: 
201-212. 
2. Svenson S (2004) Self-assembly and self-organization: Important processes — but can 
we predict them? / Dispersion Sci Technol 25:101-118. 
3. Svenson S (ed.) (2004) Carrier-based Drug Delivery Vol. 879. ACS Symposium Series, 
American Chemical Society, Washington, DC. 
4. (a) Tomalia DA (2004) Birth of a new macromolecular architecture: Dendrimers as quantized 
building blocks for nanoscale synthetic organic chemistry. Aldrichimica Acta 37: 
300 Svenson & Tomalia 
39-57; (b) Tomalia DA (2005) Birth of a new macromolecular architecture: Dendrimers 
as quantized building blocks for nanoscale synthetic polymer chemistry. Prog Polym 
Sci 30:294-324; (c) Tomalia DA (2005) The dendritic state. Materials Today March: 34-46; 
(d) Tomalia DA (2005) Dendrimeric supramolecular and supramacromolecular assemblies, 
in Supramolecular Polymers, 2nd Ed., CRC Press, Taylor & Francis, Boca Raton, FL. 
5. Tomalia DA and Frechet JMJ (eds.) (2001) Dendrimers and Other Dendritic Polymers. J. 
Wiley & Sons Ltd., Chichester. 
6. (a) Watkins DM, Sayed-Sweet Y, Klimash JW, Turro NJ and Tomalia DA (1997) 
Dendrimers with hydrophobic cores and the formation of supramolecular dendrimer 
— surfactant assemblies. Langmuir 13:3136-3141; (b) Sayed-Sweet Y, 
Hedstrand DM, Spinder R and Tomalia DA (1997) Hydrophobically modified 
poly(amidoamine) (PAMAM) dendrimers: Their properties at the air-water interface 
and use as nanoscopic container molecules. / Mater Chem 7:1199-1205. 
7. Recker J, Tomcik DJ and Parquette JR (2000) Folding dendrons: The development 
of solvent-, temperature-, and generation-dependent chiral conformational order in 
intramolecularly hydrogen-bonded dendrons. J Am Chem Soc 122:10298-10307. 
8. Boas U, Karlsson AJ, de Waal BFM and Meijer EW (2001) Synthesis and properties of 
new thiourea-functionalized poly(propylene imine) dendrimers and their role as hosts 
for urea functionalized guests. / Org Chem 66:2136-2145. 
9. Tomalia DA, Naylor AM and Goddard III WA (1990) Starburst dendrimers: Molecular 
level control of size, shape, surface chemistry, topology and flexibility from atoms to 
macroscopic matter. Angew Chem Int Ed Engl 29:138-175. 
10. Naylor AM, Goddard III WA, Kiefer GE and Tomalia DA (1989) Starburst dendrimers 5. 
Molecular shape control. } Am Chem Soc 111:2339-2341. 
11. Jansen JFGA, Debrabandervandenberg EMM and Meijer EW (1994) Encapsulation of 
guest molecules into a dendritic box. Science 266:1226-1229. 
12. Jansen JFGA, Meijer EW and Debrabandervandenberg EMM (1995) The dendritic box: 
Shape-selective liberation of encapsulated guests. J Am Chem Soc 117:4417-^1418. 
13. (a) Esfand R and Tomalia DA (2001) Poly(amidoamine) (PAMAM) dendrimers: From 
biomimicry to drug delivery and biomedical applications. Drug Disc Today 6:427-436; 
(b) Tomalia DA, Hall M, Hedstrand (1987) STARBURST® Dendrimers III. The importance 
of branch junction symmetry in the development of topological shell molecules. 
J Am Chem Soc 109:1601-1603. 
14. Twyman LJ, Beezer AE, Esfand R, Hardy MJ and Mitchell JC (1999) The synthesis of 
water-soluble dendrimers, and their application as possible drug delivery systems. 
Tetrahedron Lett 40:1743-1746. 
15. Gorman CB and Smith JC (2001) Structure-property relationships in dendritic encapsulation. 
Ace Chem Res 34:60-71. 
16. Chasse TL, Yohannan JC, Kim N, Li Q, Li Z and Gorman CB (2003) Dendritic 
encapsulation-roles of cores and branches. Tetrahedron 59:3853-3861. 
17. Chasse TL, Sachdeva R, Li Q, Li Z, Petrie RJ and Gorman CB (2003) Structural effects 
on encapsulation as probed in redox-active core dendrimer isomers. / Am Chem Soc 
125:8250-8254. 
Dendrimers as Nanoparticulate Drug Carriers 301 
18. Wallimann P, Marti T, Furer A and Diederich F (1997) Steroids in molecular recognition. 
Chem Rev 97:1567-1608. 
19. Wallimann P, Seiler P and Diederich F (1996) Dendrophanes: Novel steroid-recognizing 
dendritic receptors. Helv Chim Acta 79:779-788. 
20. Smith DK, Zingg A and Diederich F (1999) Dendroclefts: Optically active dendritic 
receptors for the selective recognition and chiroptical sensing of monosaccharide 
guests. Helv Chim Acta 82:1225-1241. 
21. Smith DK and Diederich F (1998) Dendritic hydrogen bonding receptors: Enantiomerically 
pure dendroclefts for the selective recognition of monosaccharides. Chem Commun 
2501-2502. 
22. Zimmerman SC, Wendland MS, Rakow NA, Zharov I and Suslick KS (2002) Synthetic 
hosts by monomolecular imprinting inside dendrimers. Nature 418: 399-403. 
23. Broeren MAC, van Dongen JLJ, Pittelkow M, Christensen JB, van Genderen MHP and 
Meijer EW (2004) Multivalency in the gas phase: The study of dendritic aggregates by 
mass spectrometry. Angew Chem hit Ed 43:3557-3562. 
24. Maeda H and Matsumura Y (1986) A new concept in macromolecular therapeutics in 
cancer chemotherapy: Mechanism of tumoritropic accumulation of proteins and antitumor 
agent SMANCS. Cancer Res 46:6387-9392. 
25. Boas U and Heegaard PMH (2004) Dendrimers in drug research. Chem Soc Rev 33: 
43-63. 
26. Frechet JMJ (2000) Designing dendrimers for drug delivery. Pharm Sci Technol Today 
2:393-401. 
27. Gillies ER and Frechet JMJ (2005) Dendrimers and dendritic polymers in drug delivery. 
Drug Disc Today 10:35^2. 
28. Malik N, Evagorou EG and Duncan R (1999) Dendrimer-platinate: A novel approach 
to cancer chemotherapy. Anticancer Drugs 10:767-776. 
29. (a) Malik N and Duncan R (2003) Dendritic-platinate drug delivery system. US 
6585956; (b) Malik N and Duncan R (2004) Method of treating cancerous tumors 
with a dendritic-platinate drug delivery system. US 6 790 437; (c) Malik N, Duncan R, 
TomaliaDA, Esfand R (2006) Dendritic-antineoplastic drug delivery system. US 
7005124. Patents owned by Dendritic Nanotechnologies, Inc. (DNT). 
30. Kelland LR and Farrell NP (eds.) (2000) Platinum-Based Drugs in Cancer Chemotherapy. 
Humana Press: Totowa, NJ. 
31. Gianasi E, Wasil M, Evagorou EG, Keddle A, Wilson G and Duncan R (1999) HPMA 
copolymer platinates as novel antitumor agents: In vitro properties, pharmacokinetics 
and antitumor activity in vivo. Eur J Cancer 35:994-1002. 
32. Balogh L, Swanson DR, Tomalia DA, Hagnauer GL and McManus AT (2001) Dendrimersilver 
complexes and nanocomposites as antimicrobial agents. Nano Lett 1:18-21. 
33. Kojima C, Kono K, Maruyama K and Takagishi T (2000) Synthesis of polyamidoamine 
dendrimers having poly(ethylene glycol) grafts and their ability to encapsulate anticancer 
drugs. Bioconjug Chem 11:910-917. 
34. Pan GF, Lemmouchi Y, Akala EO and Bakare O (2005) Studies on PEGylated and drugloaded 
PAMAM dendrimers. / Bioactive Compat Polym 20:113-128. 
302 Svenson & Tom all a 
35. Bhadra D, Bhadra S, Jain S and Jain NK (2003) A PEGylated dendritic nanoparticulate 
carrier of fluorouracil. Int J Pharm 257:111-124. 
36. Wang F, Bronich TK, Kabanov AV, Rauh RD and Roovers J (2005) Synthesis and evaluation 
of a star amphiphilic block copolymer from poly (e-caprolactone) and poly(ethylene 
glycol) as a potential drug delivery carrier. Bioconjug Chem 16:397-405. 
37. Namazi H and Adell M (2005) Dendrimers of citric acid and poly(ethylene glycol) as 
the new drug delivery agents. Biomaterials 26:1175-1183. 
38. Yang H and Lopina ST (2005) Extended release of a novel antidepressant, venlafaxine, 
based on anionic poly(amidoamine) dendrimers and poly(ethylene glycol)-containing 
semi-interpenetrating networks. / Biomed Mater Res Part A 72A:107-114. 
39. Kolhe P, Misra E, Kannan RM, Kannan S and Lieh-Lai M (2003) Drug complexation, in 
vitro release and cellular entry of dendrimers and hyperbranched polymers. Int J Pharm 
259:143-160. 
40. Kannan S, Kolhe P, Raykova V, Glibatec M, Kannan RM, Lieh-Lai M and Bassett D 
(2004) Dynamics of cellular entry and drug delivery by dendritic polymers into human 
lung epithelial carcinoma cells. / Biomater Sci Polym Ed 15:311-330. 
41. Chauhan AS, Jain NK, Diwan PV and Khopade AJ (2004) Solubility enhancement 
of indomethacin with poly(amidoamine) dendrimers and targeting to inflammatory 
regions of arthritic rats. / Drug Targ 12:575-583. 
42. Kubasiak LA, Chauhan AS and Tomalia DA (2005) Manuscript in preparation. 
43. Devarakonda B, Hill RA and de Villiers MM (2004) The effect of PAMAM dendrimer 
generation size and surface functional group on the aqueous solubility of nifedipine. 
Int]Pharm 284:133-140. 
44. Potluri SK, Ramulu AR and Pardhasaradhi M (2004) Synthesis of new unsymmetrical 
optically active (s)-(+)-naproxen. Tetrahedron 60:10915-10920. 
45. Ooya T, Lee J and Park K (2004) Hydrotropic dendrimers of generations 4 and 5: Synthesis, 
characterization, and hydrotropic solubilization of paclitaxel. Bioconjug Chem 
15:1221-1229. 
46. Khandare J, Kolhe P, Pillai O, Kannan S, Lieh-Lai M and Kannan RM (2005) Synthesis, 
cellular transport, and activity of poly(amidoamine) dendrimer-methylprednisolone 
conjugates. Bioconjug Chem 16:330-337. 
47. Shamis M, Lode HN and Shabat D (2004) Bioactivation of self-immolative dendritic 
prodrugs by catalytic antibody 38C2. / Am Chem Soc 126:1726-1731. 
48. Haba K, Popkov M, Shamis M, Lerner RA, Barbas III CF and Shabat D (2005) Singletriggered 
trimeric prodrugs. Angew Chem Int Ed 44:716-720. 
49. Padilla De Jesus OL, Ihre HR, Gagne L, Frechet JMJ and Szoka Jr. FC (2002) Polyester 
dendritic systems for drug delivery applications: In vitro and in vivo evaluation. 
Bioconjug Chem 13:453-461. 
50. Nishiyama N, Stapert HR, Zhang GD, Takasu D, Jiang DL, Nagano T, Aida T and 
Kataoka K (2003) Light-harvesting ionic dendrimer porphyrins as new photosensitizers 
for photodynamic therapy. Bioconjug Chem 14:58-66. 
51. Zhang GD, Harada A, Nishiyama N, Jiang DL, Koyama H, Aida T and Kataoka K 
(2003) Polyion complex micelles entrapping cationic dendrimer porphyrin: Effective 
photosensitizer for photodynamic therapy of cancer. / Control Rel 93:141-150. 
Denclrimers as Nanoparticulate Drug Carriers 303 
52. Jang WD, Nishiyama N, Zhang GD, Harada A, Jiang DL, Kawauchi S, Morimoto Y, 
Kikuchi M, Koyana H, Aida T and Kataoka K (2005) Supramolecular nanocarrier 
of anionic dendrimer porphyrins with cationic block copolymers modified with 
poly(ethylene glycol) to enhance intracellular photodynamic efficacy. Angew Chem Int 
Ed 44:419^123. 
53. Battah SH, Chee CE, akanishi H, Gerscher S, MacRobert AJ and Edwards C (2001) 
Synthesis and biological studies of 5-aminolevulinic acid-containing dendrimers for 
photodynamic therapy. Bioconjug Chem 12:980-988. 
54. Paul A, Hackbarth S, Molich A, Luban C, Oelckers S, Bohm F and Roder B (2003) Comparative 
study of the photosensitization of Jurkat cells in vitro by pheophorbide a and a 
pheophorbide a-diaminobutane poly(propylene imine) dendrimer complex. Laser Phys 
13:22-29. 
55. Wu G, Barth RF, Yang WL, Chatterjee M, Tjarks W, Ciesielski MJ and Fenstermaker RA 
(2004) Site-specific conjugation of boron-containing dendrimers to anti-EGF receptor 
monoclonal antibody cetuximab (IMC-C225) and its evaluation as a potential delivery 
agent for neutron capture therapy. Bioconjug Chem 15:185-194. 
56. Shukla S, Wu G, Chatterjee M, Yang WL, Sekido M, Diop LA, Muller R, Sudimack 
JJ, Lee RJ, Barth RF and Tjarks W (2003) Synthesis and biological evaluation of folate 
receptor-targeted boronated PAMAM dendrimers as potential agents for neutron capture 
therapy. Bioconjug Chem 14:158-167. 
57. Kono K, Liu M and Frechet JMJ (1999) Design of dendritic macromolecules containing 
folate or methotrexate residues. Bioconjug Chem 10:1115-1121. 
58. Quintana A, Raczka E, Piehler L, Lee I, Myc A, Majoros I, Patri AK, Thomas T, 
Mule J and Baker Jr. JR (2002) Design and function of a dendrimer-based therapeutic 
nanodevice targeted to tumor cells through the folate receptor. Pharm Res 
19:1310-1316. 
59. Ross JF, Chaudhuri PK and Ratnam M (1994) Differential regulation of folate receptor 
isoforms in normal and malignant tissues in vivo and established cell lines. Physiologic 
and clinical implications. Cancer 73:2432-2443. 
60. Tansey W, Ke S, Cao XY, Pasuelo MJ, Wallace S and Li C (2004) Synthesis and characterization 
of branched poly(L-glutamic acid) as a biodegradable drug carrier. / Control 
Rel 94:39-51. 
61. Lundquist JJ and Toone EJ (2002) The cluster glycoside effect. Chem Rev 102:555-578. 
62. Zanini D and Roy R (1997) Synthesis of new a-thiosialodendrimers and their binding 
properties to the sialic acid specific lectin from Limax flavus. / Am Chem Soc 119:2088- 
2095. 
63. Bezouska K, Pospisil MF, Vannucci LF, Fiserova AF, Krausova KF, Horvath OF, Kren VF, 
Mosca FF, Lindhorst TK, Sadalapure KF and Bezouska K (2002) Design, functional evaluation 
and biomedical applications of carbohydrate dendrimers (glycodendrimers). 
Rev Mol Biotechnol 90:269-290. 
64. Roy R (1996) Syntheses and some applications of chemically defined multivalent glycoconjugates. 
Curr Ovin Struct Biol 6:692-702. 
65. Lindhorst TK (2002) Artificial multivalent sugar ligands to understand and manipulate 
carbohydrate-protein interactions. Top Curr Chem Host-Guest Chem 218:201-235. 
304 Svenson & Tomalia 
66. Rockendorf N and Lindhorst TK (2001) Glycodendrimers. Top Curr Chem Dend IV 
217:201-238. 
67. Veprek P and Jezek J (1999) Peptide and glycopeptide dendrimers. Part II. / Pept Sci 
5:203-220. 
68. Andre S, Pieters RJ, Vrasidas I, Kaltner H, Kuwabara I, Liu FT, Liskamp RM and Gabius 
HJ (2001) Wedgelike glycodendrimers as inhibitors of binding of mammalian galectins 
to glycoproteins, lactose maxiclusters, and cell surface glycoconjugates. ChemBioChem 
2:822-830. 
69. Pieters RJ (2004) Interference with lectin binding and bacterial adhesion by multivalent 
carbohydrates and peptidic carbohydrate mimics. Trends Glycosci Glycotechnol 16: 
243-254. 
70. Baek MG and Roy R (2002) Synthesis and protein binding properties of T-antigen containing 
GlycoPAMAM dendrimers. Bioorg Med Chem 10:11-17. 
71. Roy R, Baek MG and Rittenhouse-Olson K (2001) Synthesis of N,N'- 
bis(acrylamido)acetic acid-based T-antigen glycodendrimers and their mouse monoclonal 
IgG antibody binding properties. / Am Chem Soc 123:1809-1816. 
72. Roy R and Baek MG (2002) Glycodendrimers: Novel glycotope isosteres unmasking 
sugar coating. Case study with T-antigen markers from breast cancer MUC1 glycoprotein. 
Rev Mol Biotechnol 90:291-309. 
73. Benito JM, Gomez-Garcia M, Mellet CO, Baussanne I, Defaye J and Fernandez JMG 
(2004) Optimizing saccharide-directed molecular delivery to biological receptors: 
Design, synthesis, and biological evaluation of glycodendrimer-cyclodextrin conjugates. 
J Am Chem Soc 126:10355-10363. 
74. Hansen HC, Haataja S, Finne J and Magnusson G (1997) Di-, tri-, and tetravalent dendritic 
galabiosides that inhibit hemagglutination by Streptococcus suis at nanomolar 
concentration. / Am Chem Soc 119:6974-6979. 
75. Rendle PM, Seger A, Rodrigues J, Oldham NJ, Bott RR, Jones JB, Cowan MM and Davies 
BG (2004) Glycodendriproteins: A synthetic glycoprotein mimic enzyme with branched 
sugar-display potently inhibits bacterial aggregation. / Am Chem Soc 126:4750^1751. 
76. Shaunak S, Thomas S, Gianasi E, Godwin A, Jones E, Teo I, Mireskandari K, Luthert P, 
Duncan R, Patterson S, Khaw P and Brocchini S (2004) Polyvalent dendrimer glucosamine 
conjugates prevent scar tissue formation. Nat Biotech 22:977-984. 
77. Thomas TP, Patri AK, Myc A, Myaing MT, Ye JY, Norris TB and Baker Jr JR (2004) In 
vitro targeting of synthesized antibody-conjugated dendrimer nanoparticles. Biomacromol 
5:2269-2274. 
78. Hong MY, Yoon HC and Kim HS (2003) Protein-ligand interactions at poly 
(amidoamine) dendrimer monolayers on gold. Langmuir 19:416-421. 
79. Sanchez-Sancho F, Perez-Inestrosa E, Suau R, Mayorga C, Torres MJ and Blanca M 
(2002) Dendrimers as carrier protein mimetics for IgE antibody recognition. Synthesis 
and characterization of densely penicilloylated dendrimers. Bioconjug Chem 13:647-653. 
80. Yang H and Lopina ST (2003) Penicillin V-conjugated PEG-PAMAM star polymers. 
/ Biomater Sci-Polym Ed 14:1043-1056. 
Dendrimers as Nanoparticulate Drug Carriers 305 
81. Bourne N, Stanberry LR, Kern ER, Holan G, Matthews B and Bernstein DI (2000) 
Dendrimers, a new class of candidate topical microbicides with activity against herpes 
simplex virus infection. Antimicrob Agents Chemother 44:2471-2474. 
82. Product Focus: VivaGel™, Starpharma Limited, Melbourne, Australia. 
83. Gong Y, Matthews B, Cheung D, Tarn T, Gadawski I, Leung D, Holan G, Raff J and 
Sacks S (2002) Evidence of dual sites of action of dendrimers: SPL-2999 inhibits both 
virus entry and late stages of herpes simplex virus replication. Antiviral Res 55:319-329. 
84. Witvrouw M, Fikkert V, Pluymers W, Matthews B, Mardel K, Schols D, Raff J, Debyser Z, 
DeClercq E, Holan G and Pannecouque C (2000) Polyanionic (i.e. polysulfonate) 
dendrimers can inhibit the replication of human immunodeficiency virus by interfering 
with both virus adsorption and later steps (Reverse transcriptase/integrase) in 
the virus replicative cycle. Mol Pharmacol 58:1100-1108. 
85. Chen CZ and Cooper SL (2002) Interactions between dendrimer biocides and bacterial 
membranes. Biomaterials 23:3359-3368. 
86. Chen CZ, Beck-Tan NC, Dhurjati P, van Dyk TK, LaRossa Ra and Cooper SL (2000) 
Quaternary ammonium functionalized poly(propylene imine) dendrimers as effective 
antimicrobials: Structure-activity studies. Biomacromol 1:473^80. 
87. Nagahori N, Lee RT, Nishimura S, Page D, Roy R and Lee YC (2002) Inhibition 
of adhesion of type 1 fimbriated Escherichia coli to highly mannosylated ligands. Chem- 
BioChem 3:836-844. 
88. Sashiwa H and Aiba SI (2004) Chemically modified chitin and chitosan as biomaterials. 
Prog Polymer Sci 29:887-908. 
89. Lebreton S, Newcombe N and Bradley M (2003) Antibacterial single-bead screening. 
Tetrahedron 59:10213-10222. 
90. Solassol J, Crozet C, Perrier V, Leclaire J, Beranger F, Caminade AM, Meunier B, 
Dormont D, Majoral JP and Lehmann S (2004) Cationic phosphorus-containing dendrimers 
reduce prion replication both in cell culture and in mice infected with scapie. 
/ Gen Virol 85:1791-1799. 
91. El-Sayed M, Rhodes CA, Ginski M and Ghandehari H (2003) Transport mechanism(s) of 
poly(amidoamine) dendrimers across Caco-2 cell monolayers. IntJ Pharm 265:151-157. 
92. D'Emanuele A, Jevprasesphant R, Penny J and Attwood D (2004) The use of a 
dendrimer-propanolol prodrug to bypass efflux transporters and enhance oral bioavailability. 
/ Control Rel 95:447-453. 
93. Chauhan AS, Sridevi S, Chalasani KB, Jain AK, Jain SK, Jain NK and Diwan PV (2003) 
Dendrimer-mediated transdermal delivery: Enhanced bioavailability of indomethacin. 
/ Control Rel 90:335-343. 
94. Wang ZX, Itoh YS, Hosaka Y, Kobayashi I, Nakano Y, Maeda I, Umeda F, Yamakawa J, 
Kawase M and Yagi K (2003) Novel transdermal drug delivery system with 
polyhydroxyalkanoate and starburst poly(amidoamine) dendrimer. / Biosci Bioeng 
95:541-543. 
95. Wang ZX, Itoh YS, Hosaka Y, Kobayashi I, Nakano Y, Maeda I, Umeda F, Yamakawa J, 
Nishimine M, Suenobu T, Fukuzumi S, Kawase M and Yagi K (2003) Mechanism of 
306 Svenson & Tomalia 
enhancement effect of dendrimer on transdermal drug permeation through polyhydroxyalkanoate 
matrix. / Biosci Bioeng 96:537-540. 
96. Vandamme TF and Brobeck L (2005) Poly(amidoamine) dendrimers as ophthalmic 
vehicles for ocular delivery of pilocarpine nitrate and tropicamide. / Control Rel 102: 
23-38. 
97. Jevprasesphant R, Penny J, Jalal R, Attwood D, McKeown NB and D'Emanuele A (2003) 
The influence of surface modification on the cytotoxicity of PAMAM dendrimers. Int J 
Pharm 252:263-266. 
98. El-Sayed M, Ginski M, Rhodes C and Ghandehari H (2002) Transepithelial transport 
of poly(amidoamine) dendrimers across Caco-2 cell monolayers. / Control Rel 
81:355-365. 
99. Fischer D, Li Y, Ahlemeyer B, Krieglstein J and Kissel T (2003) In vitro cytotoxicity 
testing of polycations: Influence of polymer structure on cell viability and hemolysis. 
Biomaterials 24:1121-1131. 
100. Malik N, Wiwattanapatapee R, Klopsch R, Lorenz K, Frey H, Weener JW, Meijer EW, 
Paulus W and Duncan R (2000) Dendrimers: Relationship between structure and biocompatibility 
in vitro, and preliminary studies on the biodistribution of I-125-labelled 
poly(amidoamine) dendrimers in vivo.} Control Rel 65:133-148. 
101. Zinselmeyer BH, Mackay SP, Schatzlein AG and Uchegbu IF (2002) The lowergeneration 
polypropylenimine dendrimers are effective gene-transfer agents. Pharm 
Res 19:960-967. 
102. Kubasiak LA and Tomalia DA (2005) Manuscript in preparation. 
103. Yoo H and Juliano RL (2000) Enhanced delivery of antisense oligonucleotides with 
fluorophore-conjugated PAMAM dendrimers. Nucleic Acids Res 28:4225-4231. 
104. Roberts JC, Bhalgat MK and Zera RT (1996) Preliminary biological evaluation of 
poly(amidoamine) (PAMAM) starburst dendrimers. / Biomed Mater Res 30:53-65. 
105. Kobayashi H, Kawamoto S, Saga T, Sato N, Hiraga A, Ishimori T, KonishiJ, Togashi K 
and Brechbiel MW (2001) Positive effects of polyethylene glycol conjugation to 
generation-4 polyamidoamine dendrimers as macromolecular MR contrast agents. 
Magn Reson Med 46:781-788. 
106. Tomalia DA and Frechet JMJ (2002) Discovery of dendrimers and dendritic polymers: 
A brief historical perspective. / Polym Sci Part A: Polym Chem 40:2719-2728. 
14 
Drug Nanocrystals/Nanosuspensions for 
the Delivery of Poorly Soluble Drugs 
Rainer H. Muller and Jens-Uwe A. H. Junghanns 
1. Introduction 
Since the last ten years, the number of poorly soluble drugs is steadily increasing. 
According to estimates, about 40% of the drugs in the pipelines have solubility 
problems.1 The increased use of high throughput screening methods leads to the 
discovery of more drugs being poorly water soluble. In the literature, figures are 
quoted that about 60 percent of the drugs coming directly from synthesis are nowadays 
poorly soluble.2 Poor solubility is not only a problem for the formulation 
development and clinical testing, it is also an obstacle at the very beginning when 
screening new compounds for pharmacological activity. From this, there is a definite 
need for smart technological formulation approaches to make such poorly 
soluble drugs bioavailable. Making such drugs bioavailable means that they show 
sufficiently high absorption after oral administration, or they can alternatively be 
injected intravenously. 
There is quite a number of formulation approaches for poorly soluble drugs 
which can be specified as "specific approaches". These approaches are suitable 
for molecules having special properties with regard to their chemistry (e.g. solubility 
in certain organic media) or to the molecular size or conformation (e.g. 
molecules to be incorporated into the cyclodextrin ring structure). Of course it 
would be much smarter to have a "universal formulation approach" applicable 
to any molecule. Such a universal formulation approach to increase the oral 
307 
308 Muller & Junghanns 
bioavailability is micronization, meaning the transfer of drug powders into the size 
range between typically 1-10 /j-va. However, nowadays many drugs are so poorly 
soluble that micronization is not sufficient. The increase in surface area, and thus 
consequently in dissolution velocity, is not sufficient to overcome the bioavailability 
problems of very poorly soluble drugs of the biopharmaceutical specification 
class II. A consequent next step was to move from micronization to nanonization. 
Since the beginning of the 90s, the company Nanosystems propagated the use of 
nanocrystals (instead of microcrystals) for oral bioavailability enhancement, and 
also to use nanocrystals suspended in water (nanosuspensions) for intravenous or 
pulmonary drug delivery. 
The solution was simple; in general, simple solutions possess the smartness that 
they can be realized easier than complex systems and introduction to the market is 
faster. Nevertheless, it took about ten years before the first nanocrystals in a tablet 
appeared on the market, the product Rapamune® by the company Wyeth in 2000. 
Compared with liposomes developed in 19683 with the first products on the market 
around 1990 (e.g. Alveofact®, a lung surfactant), this was still relatively fast. What 
were the reasons that it took about one decade for nanocrystals to enter the market? 
From our point of view, pharmaceutical companies prefer to use formulation 
technology already established with know how available in the company. In addition, 
if formulation technologies are established, a company also has the possibility 
for production of the final product. Therefore, all the traditional formulation 
approaches were exploited to solve a formulation problem. In addition, formulation 
approaches were preferred, being even simpler than nanocrystals. For example, 
production of drug-containing microemulsions administered in a capsule is, 
in many cases, even simpler. Another reason for the reluctance of pharmaceutical 
companies at the beginning was the lack of large scale production methods. These 
were not available at the very beginning of the development of the nanocrystal technology. 
Meanwhile, this has changed and the major pharmaceutical companies try 
to secure or have already secured their access to nanocrystal technology. Access 
to nanocrystal technology is possible either by licencing in or alternatively by the 
attempt to develop one's own production technologies for the nanocrystals, which 
do not depend on already existing intelectual property (IP). This chapter discusses 
the physicochemical properties of nanocrystals which make them interesting for 
drug delivery, reviews and discusses briefly the various production methods available 
and highlights the opportunities for improved drug delivery using different 
application routes. 
2. Definitions 
Drug nanocrystals are crystals with a size in the nanometer range, meaning that 
they are nanoparticles with a crystalline character. There are discussions about the 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 309 
definition of a nanoparticle, referring to the size of a particle to be classified as a 
nanoparticle. Depending on the discipline, e.g. in colloid chemistry, particles are 
only considered as nanoparticles when they are in sizes below 100 nm or even below 
20 nm. Based on the size unit, in the pharmaceutical area, nanoparticles should be 
defined as having a size between a few nanometers and 1000 nm (1 /im); thus, 
microparticles possess consequently a size 1-1000 micrometer. 
A further characteristic is that drug nanocrystals are composed of 100% drug; 
there is no carrier material as in polymeric nanoparticles. Dispersion of drug 
nanocrystals in liquid media leads to "nanosuspensions", in contrast to "microsuspensions" 
or "macrosuspensions". In general, the dispersed particles need to 
be stabilized, e.g. by surfactants or polymeric stabilizers. Dispersion media can be 
water, aqueous solutions or non-aqueous media [e.g. liquid polyethylene glycol 
(PEG), oils]. Depending on the production technology, processing of drug microcrystals 
to drug nanoparticles can lead to either a crystalline or to an amorphous 
product, especially when applying precipitation. In the strict sense, such an amorphous 
drug nanoparticle should not be called nanocrystal. However, one often 
refers to "nanocrystals in the amorphous state". 
3. Physicochemical Properties of Drug Nanocrystals 
3.1. Change of dissolution velocity 
The reason for micronization is to increase the surface area, thus consequently 
according to the Noyes-Whitney equation, increasing the dissolution velocity. 
Therefore, micronization can be succesfuUy employed if the dissolution velocity 
is the rate-limiting step for oral absorption (drugs of BSC II). Of course, by moving 
one dimension further to smaller particles, the surface area is further enlarged 
and consequently, the dissolution velocity is further enhanced. In most cases, a low 
dissolution velocity is correlated with a low saturation solubility. 
3.2. Saturation solubility 
The general textbook statement is that the saturation solubility cs is a constant 
depending on the compound, the dissolution medium and the temperature. This 
is valid for powders of daily life with a size in the micrometer range or above. 
However, below a critical size of 1-2 /zm, the saturation solubility is also a function 
of the particle size. It increases with decreasing particle size below 1000 nm. 
Therefore, drug nanocrystals possess an increased saturation solubility. This has 
two advantages: 
1. According to Noyes-Whitney, the dissolution velocity is further enhanced 
because dc/dt is proportional to the concentration gradient (cs — cx)/h (cx — 
bulk concentration, h — diffusional distance). 
310 Muller & Junghanns 
2. Due to the increased saturation solubility, the concentration gradient between 
gut lumen and blood is increased, consequently, the absorption by passive 
diffusion. 
The interesting question very often asked is "How manyfold is the increased saturation 
solubility?". Data published in the literature or available to us from discussions 
range from 2-14 fold. What are the factors affecting the increase in saturation solubility? 
The factors can be identified when looking at the theoretical background. 
The Kelvin equation describes the increase in the vapor pressure of droplets in a 
gas medium as a function of their particle size, i.e. as a function of their curvature: 
, P -y*VL*cos6 
[pj~ rK*RT 
Fig. 1. The Kelvin equation. 
P = vapor pressure 
Po = equilibrium pressure of a flat liquid surface 
y = surface tension 
VL = molar volume 
cos(#) = contact angle 
rK = radius of droplet 
R = universal gas constant 
T = absolute temperature (K) 
The vapor pressure increases with increasing curvature of the surface, that means 
decreasing particle size. Each liquid has its compound specific vapor pressure, 
thus the increase in vapor pressure will be influenced by the available compoundspecific 
vapor pressure. The situation of a transfer of molecules from a liquid phase 
(droplet) to a qas phase is in principal identical to the transfer of molecules from a 
solid phase (nanocrystal) to a liquid phase (dispersion medium). The vapor pressure 
is equivalent to the dissolution pressure. In the state of saturation solubility, 
there is an equilibrium of molecules dissolving and molecules recrystallizing. This 
equilibrium can be shifted in case the dissolution pressure increases, thus increasing 
the saturation solubility. Identical to liquids with different vapor pressures under 
normal conditions (micrometer droplet size), each drug crystal has a specific dissolution 
pressure in micrometer size. 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 311 
Relative vapor pressure at 293 K 
CD 
Droplet size [u,m] 
Fig. 2. Comparison of the relative increase in vapor pressure between water, ether and oleic 
acid (calculated using the Kelvin equation) as a function of the droplet size (with permission 
after4). 
The important question is how the dissolution pressure changes, depending on 
the specific dissolution pressure of each compound and on the particle size. Model 
calculations were performed applying the Kelvin equation to compounds with 
different vapor pressures (droplets) as a function of droplet size (Fig. 2). Liquids 
with low medium and high vapor pressure were selected, such as oleic acid as an 
oil, water and ether. The important result for a drug formulation was: 
1. The increase in vapor pressure is more pronounced for compounds having 
a priori a low vapor pressure. Applied to solid compounds, increase in dissolution 
pressure will be more pronounced for compounds having a priori a low 
dissolution pressure, i.e. the relative increase is highest for poorly soluble drugs. 
2. The increase in vapor pressure is exponential, with a very pronounced increase 
occurring at droplet sizes below 100 nm. 
Figure 3 shows a calculated increase for barium sulfate as solid model compound. 
3.3. Does size really matter? 
Transferring this to drug nanocrystals means that really smart crystals with highest 
increase in saturation solubility should have a size of e.g. 50 nm or 20-30 nm. 
From this, it can be concluded that the slogan "size matters" is correct regarding 
the increase in saturation solubility, and consequently, the increase in dissolution 
312 Muller & Junghanns 
Saturation solubility of BaS04 in water at 293 K 
1,2-, 
C^ 1,0- 
1 0,8- 
o 
en 
c 0,6- 
CO 
jjj 0,4- 
o 
c 
§ 0,2- 
+3 
J2 
K 0,0- 
- 0 , 2 - 
. BaS04 properties: 
\ M = 233.40 g/mol 
\ p = 4.50 g/cm3 
\ cs = 2.22 mg/L 
\ o = 26.7mN/m 
-> i- i i i •-! n - ' i i i . . . i i i . | • i 1 '—'— I ' '— ' '—' ' ' ' 111 
0,1 1 10 100 
Drug size [jam] 
Fig. 3. Increase in saturation solubility of BaS04 in water as a function of the particle size 
calculated using the Kelvin equation (with permission after4). 
velocity caused by a higher cs. It needs to be kept in mind which blood profile is 
anticipated with a certain drug. In many cases, too fast a dissolution is not desired 
(creation of high plasma peaks, reduction of tmax). There is the request to combine 
drug nanocrystals with traditional controlled release technology (e.g. coated pellets) 
to avoid too fast a dissolution, too high plasma peaks, too early a tmax and to 
reach prolonged blood levels. To summarize, the optimal drug nanocrystal size will 
depend on: 
1. Required blood profile 
2. Administration route 
In the case of i.v. injected nanocrystals, the size should be as small as possible in case 
the pharmacokinetics of a solution should be mimicked. In the event that a targeting 
is the aim (e.g. to the brain by PathFinder technology,5 the drug nanocrystals should 
possess a certain size to delay dissolution and to give them the chance to reach the 
blood-brain barrier (BBB) for internalization by the endothelial cells of the BBB.6 
3.4. Effect of amorphous particle state 
It is well known that amorphous drugs possess a higher saturation solubility, 
compared with crystalline drug material. A classical example from the literature 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 313 
is chloramphenicol palmitate. The polymorphic modification I has a solubility of 
0.13, the high energy modification II a solubility of 0.43 and the amorphous material 
of 1.6mg/mL.7'8 The same is valid for drug nanoparticles, amorphous drug 
nanoparticles possess a higher saturation solubility, compared with equally sized 
drug nanocrystals in the crystalline state. Therefore, to reach highest saturation 
solubility increase, a combination of nanometer size and amorphous state is ideal. 
However, prerequisite for exploitation in pharmaceutical products is that the amorphous 
state can be maintained for the shelf life of the product. 
4. Production Methods 
4.1. Precipitation methods 
4.1.1. Hydrosols 
The Hydrosol technology was developed by Sucker and the intellectual property 
owned by the company Sandoz, now known as Novartis.9,10 It is basically a classical 
precipitation process known to pharmacists under the term "via humida paratum" 
(v.h.p.). This v.h.p. process was employed to prepare ointments containing finely 
dispersed, precipitated drugs. The drug is dissolved in a solvent, the solvent added 
to a non-solvent leading to the precipitation of finely dispersed drug nanocrystals. 
A problem associated with this technology is that the formed nanoparticles need to 
be stabilized to avoid growth in micrometer crystals. In addition, the drug needs 
to be soluble at least in one solvent. This creates problems for the newly synthesized 
or discovered drugs, being poorly soluble in water and simultaneously in 
organic media. Lyophilization is recommended to preserve the particle size.1 To 
our knowledge, this technology has not been applied to a product to date. 
4.1.2. Amorphous drug nanoparticles (NanoMorph®) 
Depending on the precipitation methodology, drug nanoparticles can be generated 
which are in the amorphous state. A nice example are carotine nanoparticles in food 
industry.11 
A solution of the carotinoid, together with a surfactant and a digestible oil, are 
admixed into an appropriate solvent at a specific temperature. The solution is mixed 
with a protective colloid. This tranforms the hydrophilic solvent components into 
the water phase and the hydrophobic phase of the carotinoid forms a monodisperse 
phase. X-ray analysis after subsequent lyophilization shows that approximately 
90% of the carotinoid is in the amorphous state.11 
Amorphous precipitation technology is used by the company Soliqs and the 
technology is advertised under the tradename NanoMorph®. The preservation of 
314 Miiller & Junghanns 
the amorphous state could be achieved successfully for food products. To exploit 
the amorphous technology for pharmaceutical products, the stricter requirements 
for pharmaceuticals need to be met. 
4.2. Homogenization methods 
4.2A. Microfluidizertechnology 
The previous Canadian company RTP (Montreal, now Skyepharma Canada Inc.) 
employed the microfluidizer to homogenize drug suspensions. The microfluidizer 
is a jet stream homogenizer of two fluid streams collied frontally with high velocity 
(up to 1000m/sec)12 under pressures up to 4000 bar. There is a turbulant flow, 
high shear forces, particles collied leading to particle diminution to the nanometer 
range.13-15 The high pressure applied and the high streaming velocity of the lipid 
can also lead to cavitation additionally, contributing to size diminution. The patent 
describes examples requiring up to 50 passes through the microfluidizer to obtain 
a nanosuspension.16 Sometimes, up to 100 cycles are required when applying the 
microfluidizer technology. This does not pose any problem on the small lab scale, 
but it is not production friendly for larger lab scale. The dispersion medium is water. 
4.2.2. Piston-gap homogenization in water (Dissocubes®) 
In 1994, Mueller et al.17-18 developed a high pressure homogenization method 
based on piston-gap homogenizers for drug nanosuspension production. Dispersion 
medium of the suspensions was water. A piston in a large bore cylinder creates 
pressure up to 2000 bar. The suspension is pressed through a very narrow 
ring gap. The gap width is typically in the range of 3-15 micrometer at pressures 
between 1500-150 bar. There is a high streaming velocity in the gap according to the 
Bernouli equation.19 Due to the reduction in diameter from the large bore cylinder 
(e.g. 3 cm) to the homogenization gap, the dynamic pressure (streaming velocity) 
increases and simultaneously decreases the static pressure on the liquid. The liquid 
starts boiling, and gas bubbles occur which subsequently implode, when the suspension 
leaves the gap and is again under normal pressure (cavitation). Gas bubble 
formation and implosion lead to shock waves which cause particle diminution. 
The patent describes cavitation as the reason for the achieved size diminution.17,20 
Piston-gap homogenizers which can be used for the production of nanosuspensions 
are e.g. from the companies APV Gaulin, Avestin or Niro Soavi. The technology was 
aquired by Skyepharma PLC at the end of the 90s and employed in its formulation 
development.21-23 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 315 
4.2.3. Nanopure technology 
For oral administration, the drug nanosuspensions themselves are, in most cases, 
not the final products. For patient's convenience, the drug nanocrystals should be 
incorporated in traditional dry dosage form, e.g. tablets, pellets and capsules. An 
elegant method to obtain a final formulation directly is the production of nanocrystals 
in non-aqueous homogenization media. Drug nanocrystals dispersed in liquid 
polyethylene glycol (PEG) or oils can be directly filled as drug suspensions into gelatine 
or HPMC capsules. The non-aqueous homogenization technology was established 
against the teaching that cavitation is the major diminution force in high 
pressure homogenization. Efficient particle diminution could also be obtained in 
non-aqueous media.24-30 
To prepare tablets or pellets, the dispersion medium of the nanosuspension 
needs to be removed, i.e. in general, evaporated. Evaporation is faster and possible 
under milder conditions when mixtures of water with water miscible liquids are 
used, e.g. water-ethanol. To obtain isotonic nanosuspensions for intravenous injection, 
it is beneficial to homogenize in water-glycerol mixtures. The IP owned by 
Pharmasol covers, therefore, water-free dispersion media (e.g. PEG, oils) and also 
water mixtures. 
4.3. Combination Technologies 
4.3.1. Microprecipitation™ and High Shear Forces (NANOEDGE™) 
The Nanoedge technology by the company Baxter covers a combination of precipitation 
and subsequent application of high energy shear forces, preferentially high 
pressure homogenization with piston-gap homogenizers.31 As outlined in Sec. 4.1.1, 
the precipitated particles have a tendancy to grow. According to the patent by Kipp 
et ah, treatment of a precipitated suspension with energy (e.g. high shear forces) 
avoids particle growth in precipitated suspensions (= annealing process). The 
relative complex patent description can be summarized in a simplified way that 
the subsequent annealing stabilizes the obtained particle size by precipitation. As 
described in Sec. 4.1.2, precipitated particles can be amorphous or partially amorphous. 
This implies the risk that during the shelf life of a product, the amorphous 
particles can recrystalize, leading subsequently to a reduction in oral bioavailability 
or a change in pharmacokinetics after intravenous injection. The annealing process 
by Baxter converts amorphous or partially amorphous particles to completely crystalline 
material.31 
31 6 Muller & Junghanns 
4.3.2. Nanopure® XP technology 
An important criteria for a technology is its scaling up ability and the possibility 
to produce on large scale, applying "normal" production conditions. The number 
of 50-100 passes through a homogenizer as partially required for the microfluidizer 
technology16 is not production friendly. Piston-gap homogenizers (Sec. 4.2.2) 
proved to be more efficient, typically between 10-20 homogenization cycles are 
sufficient to obtain a nanosuspension. However, it would of course be desirable 
to apply even less homogenization cycles, reducing production time, potential 
product contamination by wearing of the machine and production costs. Pharmasol 
developed a new combination process, Nanopure XP (Xtended Performance)32 
leading to: 
1. Identical particle sizes compared with high pressure homogenization in water 
(Sec. 4.2.2), but at half the cycle numbers or less. 
2. Lower particle sizes at identical cycle numbers. 
The process is again a combination technology, a pre-treatment step is followd 
by a high pressure homogenization step, typically performed with a piston-gap 
homogenizer.34'35 The code for this homogenization technology is H42. Figure 4 
[Mm] 
LD - volume size distribution 
• 50% 
• 90% 
H99% 
cycle 15 
new (H42) 
cycle 40 
old technology 
Fig. 4. Comparison of the old homogenization technology (homogenization in water, 
piston-gap homogenizer) on the right side to the new technology on the left side, presented 
are the laser diffractometry (LD) diameters 50%, 90% and 99% (volume distribution, Coulter 
LS230, Beckman-Coulter/Germany) (with permission after33). 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 31 7 
demonstrates the efficiency of method processing a very hard drug material. Applying 
the novel H42 technology leads to distinctly smaller crystals after just 15 cycles, 
compared with the "old" technology of applying 40 cycles. 
5. Application Routes and Final Formulations 
5.1. Oral administra tion 
Most attractive regarding regulatory and commercial aspects is the oral administration 
route. Compared with parenteral administration, the regulatory hurdles are 
much lower. In addition, the patient prefers an oral dosage form, that is why oral 
products possess the largest percentage of the pharmaceutical market. However, 
for the oral administration route, it is generally necessary to transfer the liquid 
nanosuspension into a solid dosage form. 
Aqueous nanosuspensions can be used as a granulation fluid for producing 
tablets or as a wetting agent for pellet production. In addition, spray drying can be 
performed in order to obtain a product which can subsequently be processed to oral 
products. The first nanosuspension product in the market was Rapamune®, introduced 
in 2000 by the company Wyeth. Rapamune® is available on the market as 
oral solution, and alternatively as tablet. The tablet is more user-friendly. Comparing 
the oral bioavailabilities of solution and nanocrystal tablet, the bioavailability 
of the nanocrystals is 21 % higher compared with the solution. The oral single dose 
of Rapamune® is 1 or 2 mg, the total tablet weight being 364 mg for 1 mg formulation 
and 372 mg for the 2 mg formulation, meaning that it contains a very low 
percentage of its total weight as nanocrystals. An important point is that the drug 
nanocrystals are released from the tablet as ultrafine suspension. In the event that 
crystal aggregation takes place to a pronounced extent, the dissolution velocity, and 
subsequently, the oral bioavailability of the BSCII drugs will be reduced. Therefore, 
there is an upper limit to load tablets with nanocrystals. In case the limit is exceeded 
and nanocrystals get in contact with each other within the excipient mixture of the 
tablet, the nanocrystals might fuse to larger crystals under the compression pressure 
during tablet production. For drugs with a low oral single dose such as Sirolimus 
in Rapamune®, incorporation into tablets causes little issues. A total nanoparticle 
load of less than 1% is well below the percentage being critical.36 
The second product on the market was Emend®, introduced in 2001 by the 
Company Merck. The drug Aprepiptant is for the treatment of emesis (single dose is 
either 80 or 125 mg). Aprepiptant will only be absorbed in the upper gastrointestinal 
tract. Bearing this in mind, nanoparticles proved to be ideal in overcoming this narrow 
absorption window. The large increase in surface area due to nanonization leads 
to rapid in vivo dissolution, fast absorption and increased bioavailability.37,38 The 
formulation of a tablet from micronized bulk powder made higher doses necessary, 
31 8 Muller & Junghanns 
leading to increased side effects.39 The drug nanocrystals are contained within the 
hard gelatin capsules as pellets. Aprepiptant was formulated as capsules for it to be 
user friendly by healthcare providers and patients, and on the other hand, to make 
it applicable as pellets via a stomach tube. Currently, studies are being undertaken 
to evaluate the change in pharmacokinectics (if any) between the pellets and the 
capsules. 
All nanocrystals in these first two products were produced using the pearl 
mill technology by Nanosystems/Elan. The prerequiste was the bioavailability of 
sufficient large scale production facilities for the respective product. In general, 
the candidates of first choice for nanosuspension technology are drugs with a relatively 
low dose. It is interesting that drugs such as Naproxen are formulated as 
nanosuspension (e.g. for fast action onset and reduced gastric irritancy),40 however, 
it requires more sophisticated formulation technology to ensure the release of the 
drug nanocrystals as fine suspension when incorporated in a tablet in a relatively 
high concentration of a single dose of 250 mg. The tablet size (weight) has to be 
acceptable for the patient and that a dosing with two tablets should be avoided, for 
reasons of patient's compliance and marketing purposes. 
Alternatively, to aqueous nanosuspensions, nanosuspensions in nonaqueous 
media can be produced by the Nanopure technology (Pharmasol). Nanocrystals 
dispersed in liquid PEG or oil can be directly filled into gelatine or HPMC capsules.25 
It saves the step of water removal and subsequent dispersion of the powder in a 
liquid capsule filling medium. 
The Nanopure technology also allows production of nanocrystals in melted 
PEG (at 60° C). After solidification of the PEG nanosuspension, the drug nanocrystals 
are fixed (and kept seperated) in the solid PEG matrix. The solidified drug 
nanocrystal containing PEG can either be milled into powders and filled into the 
capsules, or alternatively, the hot liquid PEG nanosuspension can be directly filled 
into the capsules (Fig. 5, upper). 
Instead of using aqueous nanosuspensions as fluids for the wet granulation 
process or extrusion of pellet mass, the nanosuspensions can be converted into a 
dry powder which is subsequently further processed into a tablet or a capsule. It 
also appears attractive to package such powders in sachets for redispersion in water 
or soft drinks prior to oral administration. Spray-drying is the only feasable costeffective 
way to produce such powders. An attractive approach is the production 
of so-called "compounds" as described in the direct compress technology.41 The 
term "compound" does not mean a chemical compound; in powder technology, 
"compounds" are defined as freely flowable granulate powders. In the direct compress 
technology, water-insoluble polymeric particles (e.g. Eudragit RSPO, ethylcellulose) 
are dispersed in the aqueous drug suspension, and lactose is dissolved. 
The mixture is a freely flowable compound yielded by spray-drying. The lactose 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 319 
Fig. 5. Gelatin capsules filled directly with hot liquid PEG nanosuspension, solidification 
takes place in the capsules (upper) or filling of the capsules with milled solidified PEG 
nanosuspension (lower). (From Ref. 36 with permissions.) 
is responsible for the good flowing properties. The water-insoluble polymeric 
particles also contribute to the formation of flowable granules, while at the same 
time allowing the compound to be compressed in a direct compaction process into 
tablets. The polymers form the matrix structure of the tablet. Depending on the percentage 
of the insoluble polymeric particles added, the resulting tablets may disintegrate 
fast or present a prolonged release system. A prolonged release of dissolving 
nanocrystal is desired in the case of high plasma that peaks at very early times (short 
tmax) and a targeted sustained blood level. Alternatively, the drug nanocrystal compound 
can be filled into hardgelatine capsules. Due to the presence of lactose and 
surfactant from the original nanosuspension, the compounds disperse relatively 
fast in liquids. Figure 6 shows the dispersion process of a compound after layering 
it on the surface of water in a beaker. As outlined above, efficient release and redispersion 
of the drug nanocrystals in a fine, nonaggregated state is a prerequisite for 
benefiting fully from the drug nanocrystal features.42 
5.2. Parenteral administration 
Intravenous administration is the second frequently investigated route. The 
company Baxter, with its technology NANOEDGE™, is presently focusing on 
intravenous nanosuspensions. They investigated Itraconazole nanosuspensions 
intensive.44 It could be nicely shown that the side effects of the commercial product 
Sporanox® could be distinctly reduced by the administration of a nanosuspension. 
The nephrotoxicity of Sporanox® is not caused by the drug, but by the excipient 
320 Miiller &/unghanns 
0 sec 15 sec 30 sec 60 sec 120 sec 
Fig. 6. Dispersion of a drug nanocrystal compound as a function of time after layering it 
on the surface of water in a beaker (with permission after43) (Compound: Aquacoat 40%, 
Lactose, 60%.) 
used for solubilizing the drug, the hydroxypropyl-^-cyclodextrin.45'46 The itraconzole 
nanosuspension was stabilized with Tween 80 surfactant being well tolerated 
intravenously.44 
Administration of nanosuspsensions into body cavities is also of great interest, 
e.g. to increase the tolerability of the drug, to achieve a local treatment or to have a 
depot with slow release (e.g. into the blood). It could be shown that intraperitonal 
administration of a nanosuspension was well tolerated, whereas administration of a 
macrosuspension leads to irritancy [azodicarbonamide (ADA), unpublished data]. 
Intraperitonal administration can be used for local treatment or to obtain a depot 
with prolonged release into the blood. Interesting therapeutic targets include local 
inflammations, e.g. in joints. For instance, arthritic joint inflammations are caused 
by secretion products of activated macrophages. An interesting approach is therefore 
the administration of a corticoid nanosuspension directly into the joint capsule. 
The drug particles will be phagocytosed, the drug dissolves and reduces the hyperactivity 
of the macrophages. This concept is not new, being adopted by the company 
Boots in the 80s in an attempt to incorporate the corticoid prednisolone into polymeric 
nanoparticles made from PLA-GA-copolymer.47 The particle load (polymer 
load) required to achieve a therapeutic drug level was being calculated. However, 
incubating macrophage cell cultures with the required particle concentration lead 
to cytotoxicity. The concept could not be realized, as it cannot occur with drug 
nanocrystals since no carrier polymer to required and present. 
Producing parenteral products with drug nanocrystals has to meet higher 
regulatory hurdles and product quality standards distinctly. The produced drug 
nanosuspensions need to be terminally sterilized or alternatively produced in an 
aseptic process. In principal, sterilization is possible by autoclaving. However, the 
increase in temperature can reduce hydration of steric stabilizers, thus leading to 
some aggregation during the sterilization process. Gamma irradiation is a priori a 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 321 
non-preferred process by the industry due to the necessary analytics (i.e. proof of 
absence of toxic irradiation products). In addition, it was also observed that irradiation 
can cause aggregation not by directly interacting with the drug nanocrystals, 
but with the stabilizing surfactant. Irradiation of Tarazepide nanosuspension leads 
to aggregation; simultaneously, a decrease in zeta potential also occurred during 
the irradiation process. A decrease in zeta potential, i.e. electrostatic repulsion, was 
considered as the cause for the aggregation process. It can be concluded that the 
production of drug nanosuspensions in an aseptic, controlled process has to be 
preferred, compared with the terminal sterilization by irradiation. The aseptic production 
process can be validated and documented relatively easy, therefore, being 
simpler to handle as an irradiation sterilization with accompanied analytics. 
5.3. Miscellaneous administration routes 
Oral and parenteral/intravenous routes are the ones in which developments are 
focusing, clearly due to the commercial background and the relation between the 
development costs for a market product versus its potential annual sales. However, 
drug delivery could also be improved when using drug nanocrystals for pulmonary 
and ophthalmic adminstration or dermal application. 
Poorly soluble drugs could be inhaled as drug nanosuspension. The drug 
nanosuspension can be nebulized using commercially available nebulizers.48,49 Disposition 
in the lungs can be controlled via the size distribution of the generated 
aerosol droplets. Compared with microcrystals, the drug is more evenly distributed 
in the droplets when using a nanosuspension. The number of crystals are higher, 
consequently, the possibility that one or more drug crystals are present in each 
droplet is higher. 
It could be shown that nanoparticles possess a prolonged retention time in the 
eye, most likely due to their adhesive properties.50-52 From this, poorly soluble 
drugs could be administered as a nanosuspension. However, the major obstacles 
are the commercial considerations. In many cases, the sales volume do not justify 
the costs for the development of a new market product. This is especially the case 
when a company has already a drug formulation which might be less efficient, but 
is already a product on the market. The price achievable with an improved product 
is not sufficiently high to cover the development costs of this new product. An 
additional major obstacle for the development of such improved products is the cost 
reduction policy of the healthcare systems worldwide. A longer treatment time with 
a less efficient product might still be less expensive for the healthcare system than a 
shorter treatment time with a more efficient, but distinctly more expensive product. 
The same is valid for dermal products. Sales per product are lower compared 
with e.g. oral products, as the dermal market is smaller. Dermal nanosuspensions 
322 MiJller&Junghanns 
are mainly of interest if conventional formulation technology fails or if it is distinctly 
less efficient. Dermal drug nanosuspensions lead to a supersaturated system 
because of their increased saturation solubility. The higher concentration gradient 
between topical formulation and skin can improve drug penetration into the skin. 
In addition, because of their small size, drug nanocrystals could target the hair follicle 
by protruding into the gap around the hairs. This was illustrated in solid lipid 
nanoparticles of a similar size.53 Adhesive properties of drug nanocrystals are also 
an area of interest. Adherence to the skin reduces the "loss" of drug to the environment/
third persons. This is especially so in the event that highly active compounds 
are applied, e.g. hormones. For this reason, the drug estradiole was incorporated 
into solid lipid nanoparticles to better localize it on the skin.54 
6. Nanosuspensions as Intermediate Products 
As described above, nanosuspensions can be produced such that nanocrystals 
appear in final products. Alternatively, drug nanosuspensions can be used as intermediate 
product, i.e. the drug nanocrystals do not appear in the final product. 
Recently, the SolEmuls® technology was developed to produce drug-loaded emulsions 
for intravenous injection, i.e. localizing poorly soluble drugs in the interfacial 
layer of lecithin emulsions.55-57 The applicability of the technology has been proven 
for several drugs including amphotericin B,58 itraconazole,59,60 ketoconazole,61 
and carbamazepine,62'63 among others. The drug Amphotericin B is on the market 
as a solution (Fungizone®), but also in liposomes (Ambisome®); the latter 
having the benefit of reduced nephrotoxicity.64 Liposomes are relatively expensive 
(daily treatment costs approximately EUR 1000-200064,65), therefore Amphotericin 
B was incorporated into parenteral emulsions. These emulsions can also 
reduce nephrotoxicity,66 but for their production, it was necessary to use organic 
solvents. Egg lecithin and amphotericin B were dissolved in an organic solvent, 
the solvent evaporated and the obtained drug-lecithin mixture was used to produce 
an o /w emulsion. In these emulsions, amphotericin B was located in the 
interfacial lecithin layer as Amphotericin B is simultaneously poorly soluble in 
water and in oils.67 There were also attempts to incorporate amphotericin B in the 
emulsion by simply adding Amphotericin B powder to the emulsion and subsequently 
shaking it. However, even shaking for 18 hours with 1800rph was unable 
to completely dissolve the Amphotericin B. The reason was simply due to its low 
solubility in the water, and the dissolution velocity was also extremely low, i.e. 
the process of dissolution and redistribution into the lecithin layer takes too long 
for it to be used in pharmaceutical production. The problem was solved by the 
SolEmuls technology, i.e. simple co-homogenization of oil droplets and microcrystals. 
For a de novo production, a coarse pre-emulsion of lecithin stabilized 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 323 
oil droplets in water is prepared, the drug powder is admixed under stirring 
and the obtained hybrid suspension subsequently homogenized at 600 bar (pressure 
being in the range to be used in pharmaceutical production lines). The high 
streaming velocities in the homogenization process lead to fast dissolution of 
the drug microcrystals and the re-distribution into the interfacial lecithin layer 
(Fig. 7). 
Depending on the size of the drug crystals, 5-10 homogenization cycles are 
required. The number of homogenization cycles can be reduced when adding the 
. drug not as microcrystals, but as nanocrystals in the form of a nanosuspension. A 
concentrated nanosuspension is prepared (e.g. 20-30% solid content) and added to 
the pre-emulsion. Ideally the nanosuspension is also stabilized by lecithin, i.e. the 
same emulsifier for the suspension and the emulsion. Alternatively, intravenously 
accepted stabilizers such as Tween 80 or Poloxamer 188 can be used. They are 
accepted intravenously without posing any regulatory issues. In addition, mixing 
the emulsion and nanosuspension at a ratio of 10:1 or higher will dilute the stabilizer 
concentration used in the nanosuspension by at least a factor of 10, meaning that 
in the final product, the nanosuspension surfactant concentration is typically 0.1 
or 0.01%. The question might arise as to why an emulsion should be prepared 
using a nanosuspension as an intermediate product, when it can administer the 
nanosuspension itself intravenously? 
simple 
shaking 
r ~\ 
iecithin / drug mixture 
evaporation 
O rf-Q 
O ' o T 
direct production 
through highpressure 
homogenization 
dc/dt-co 
organic solution t lecithin + drug 
drug crystal or 
suspension 
Fig. 7. Drug incorporation through various methods in comparison. Left: traditional 
attempt of shaking or alternatively use of organic solvent; Right: SolEmuls® process. 
324 Muller & Junghanns 
The reason is that drug-loaded parenteral emulsions are already products on the 
market (e.g. Diazepam-Lipuro, Etomidate-Lipuro, etc.), i.e. in a dosage form with 
which the regulatory authorities are already familiar with. Applying the SolEmuls 
technology and using lecithin-stabilized nanosuspension, the final product will 
only contain the excipients of an o /w emulsion for parenteral nutrition, without 
additional excipient plus the drug. It is an accepted known system with regard 
to the excipient status and its perfomance after intravenous injection. In contrast, 
drug nanosuspensions represent a new dosage form not yet present as intravenous 
formulations on the market. Registration of a completely new dosage form for a 
certain administration route is just more complicated and timely than registration 
of a product based on an established, known technology. 
7. Perspectives 
There was a "delayed" acceptance of the nanocrystal technology in the 90s. Pharmaceutical 
companies tried to solve their formulation problems with the traditional 
formulation approaches. However, the increasing number of drugs having a very 
low solubility, and not able to be formulated with these traditional formulation 
approaches, lead to a broad acceptance of the drug nanocrystal technology. This is 
clearly reflected in the increasing number of licensing agreements between companies 
holding nanocrystal IP and a number of medium and large pharmaceutical 
companies. The smartness of the technology is that it can be universally applied 
to practically any drug. Identical to micronization, it is a universal formulation 
principle, but limited to BSC drugs class II. The time between the beginning of 
intensive research in the drug nanocrystal technology and the first products on the 
market was relatively short, about one decade. The value of a formulation principle 
or technology can be clearly judged by looking at the number of products on the 
market, in the clinical phases, and/or the time of entry into the market. Based on 
these criteria, the drug nanocrystal technology is a successful emerging technology. 
Meanwhile, "Big Pharma" also realized the drug nanocrystal value. In combination 
with the further increasing number of poorly soluble drugs, a distinct increase in 
drug nanocrystal-based products on the market can be expected. In many cases, oral 
products will dominate because of the market share, higher sales volumes and less 
regulatory hurdles and quality requirements, compared with parenteral products. 
References 
1. Speiser PP (1998) Poorly soluble drugs, a challenge in drug delivery, in Muller RH, 
Benita S and Bohm B (eds.), Emulsions and Nanosuspensions for the Formulation of Poorly 
Soluble Drugs, Medpharm Scientific Publishers: Stuttgart. 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 325 
2. Merisko-Liversidge E (2002) Nanocrystals: Resolving Pharmaceutical Formulation Issues 
Associated with Poorly Water-soluble Compounds in Particles, Marcel Dekker: Orlando. 
3. Bangham AD and Haydon DA (1968) infrastructure of membranes: Biomolecular organization. 
Brit Med Bull 24(2):124-6. 
4. Anger S (2005) PhD Thesis, in PhD Thesis Pharmaceutical Technology. Freie Universitat: 
Berlin. 
5. Miiller RH, Luck M and Kreuter J (1998) Arzneistofftragerpartikel fur die gewebsspezifische 
Arzneistoffapplikation. Europaische Patentschrift, PCT/EP98/06429. 
6. Kreuter J, Kharkevich RN and Ivanov DA (1995) Passage of peptides through the bloodbrain 
barrier with colloidal polymer particles (nanoparticles). Brain Res 674(1):171-174. 
7. Gu Chong-Hui GDJW (2001) Estimating the relative stability of polymorphs and 
hydrates from heats of solution and solubility data. / Pharm Sci 90(9):1277-1287. 
8. Hancock C and Bruno P (2000) What is the True Solubility Advantage for Amorphous 
Pharmaceuticals? Pharm Res 17(4):397-404. 
9. Gassmann P, List M, Schweitzer A and Sucker H (1994) Hydrosols — Alternatives for the 
Parenteral Applikation of Poorly Water Soluble Drugs. Eur ] Pharm Biopharm 40:64-72. 
10. List M and Sucker H (1988) Pat.No. GB 2200048. 
11. Auweter HB, Haberkorn C, Horn H, Lueddecke D and Rauschenberger V, Patent No. 
DE19637517A1. 
12. Tunick MHVH, Diane L, Cooke PH and Malin EL (2002) Transmission electron 
microscopy of Mozzarella cheeses made from microfluidized milk. / Agri Food Chem 
50(1):99-103. 
13. Bruno RPM (1999) Microfluidizer Processor Technology for High Performance Particle Size 
Reduction, Mixing and Dispersion. Microfluidizer Processor Technology. 
14. Sunstrom JEM and Marshik-Guerts B (1996) General Route to Nanocrystalline Oxids by 
Hydrodynamic Cavitation. Chem. Mater. 
15. Gruverman IJ and Thum JR, Production of Nanostructures Under Turbulent Collision Reaction 
Conditions — Application to Catalysts, Superconductors, CMP Abrasives, Ceramics and 
other Nanoparticles. Microfluidics Research. 
16. Dearns R (2000), Atovaquone pharmaceutical compositions. US patent US 6 018 080. 
17. Miiller RH, Peters K, Becker R and Kruss B (1995) Nanosuspensions — A Novel Formulation 
for the i.v. Administration of Poorly Soluble Drugs, in 1st World Meeting of the International 
Meeting on Pharmaceutics, Biopharmaceutics and Pharmaceutical Technology hosted by 
APGI/APV. Budapest. 
18. Miiller RH, Becker R, Kruss B and Peters K (1999) Pharmaceutical Nanosuspensions for 
Medicament Administration as Systems with Increased Saturation Solubility and Rate of Solution, 
in United States Patent 5,858,410. USA. 
19. Miiller RH, Jacobs C and Kayser O (2000) Nanosuspensions for the formulation of poorly 
soluble drugs, in Nielloud F and Marti-Mestres G (eds.) Pharmaceutical Emulsions and 
Suspensions, Marcel Dekker. 
20. Miiller RH, Becker R, Kruss B and Peters K (1994) Pharmazeutische Nanosuspensionen 
zur Arzneistoffapplikation als Systeme mit erhohter Sattigungsloslichkeit und Losungsgeschwindigkeit. 
German patent 4440337.2, US Patent 5.858.410 (1999). 
326 Muller & Junghanns 
21. Muller RH, Dingier A, Schneppe T and Gohla S (2000) Large scale production of solid 
lipid nanoparticles (SLNTM) and nanosuspensions (DissoCubes™). in Wise D (ed.), 
Handbook of Pharmaceutical Controlled Release Technology. 
22. Muller RH, Jacobs C and Kayser O (2003) DissoCubes — A novel formulation for poorly soluble 
and poorly bioavailable drugs, in Rathbone MJ, Hadgraft}, Roberts MS (eds.), Modified- 
Release Drug Delivery Systems, Marcel Dekker. 
23. Rabinow BE (2004) Nanosuspensions in Drug Delivery. Nat Rev 3:785-796. 
24. Muller RH (2002) Nanopure Technology for the Production of Drug Nanocrystals and Polymeric 
Particles, in 4th World Meeting ADRITELF/APV/APGI. Florence. 
25. Bushrab NF and Muller RH (2003) Nanocrystals of Poorly Soluble Drags for Oral Administration. 
New Drugs 5: 20-22. 
26. Radtke M (2001) Nanopure™ pure drug nanoparticles for the formulation of poorly 
soluble Drugs. New Drugs 3: 62-68. 
27. Fichera MA, Keck CM and Muller RH (2004) Nanopure Technology — Drug Nanocrystals 
for the Delivery of Poorly Soluble Drugs, in Particles. Orlando. 
28. Fichera MA, Wissing SA and Muller RH (2004) Effect of 4000 Bar Homogenisation Pressure 
on Particle Diminution in Drug Suspensions, in APV. Niirnberg. 
29. Keck CM, Bushrab NF and Muller RH (2004) Nanopure® Nanocrystals for Oral Delivery of 
Poorly Soluble Drugs, in Particles. Orlando. 
30. Muller RH, Mader K and Krause K (2000) Verfahren zur schonenden Herstellung von 
hochfeinen Micro-/Nanopartikeln, in PCT Application PCT/EP00/06535: Germany. 
31. Kipp JE, Wong JCT, Doty MJ and Rebbeck CL (2003) Microprecipitation Method For Preparing 
Submicron Suspensions, in United States Patent 6,607,784. Baxter International Inc. 
(Deerfield, IL): USA. 
32. Moschwitzer J and Muller RH (2005) Method for the Production of Ultrafine Submicron 
Nanosuspensions (Pat. Application). 
33. Moschwitzer J (2005) PhD Thesis in Preparation, in PhD Thesis Pharmaceutical Technology. 
Freie Universitat: Berlin. 
34. Moschwitzer J and Muller RH (2005) Effective production of ibuprofen drug nanocrystals by 
high pressure homogenization using new two-step process, in AAPS. submitted. Nashville. 
35. Moschwitzer J and Muller RH (2005) Development of a new two-step process for the effective 
production drug nanocrystals by high pressure homogenization. in AAPS. submitted. 
Nashville. 
36. Bushrab NF (2005) PhD Thesis in preparation, in PhD Thesis Pharmaceutical Technology. 
Freie Universitat: Berlin. 
37. Moschwitzer J and Muller RH (2004) From the Drug Nanocrystal to the Final Mucoadhesive 
oral Dosage Form, in International Meeting on Pharmaceutics, Biopharmaceutics & Pharmaceutical 
Technology. Niirnberg. 
38. Moschwitzer J and Muller RH (2004) Nanosuspensions as Formulation Principle for Chemical 
Stabilization of Chemically Labile Drugs, in International Meeting on Pharmaceutics, Biopharmaceutics 
& Pharmaceutical Technology. Niirnberg. 
39. Wua YL, Landisb A, Hettricka E, Novaka L, Lynna L, Chenc K, Thompson A, Higgins R, 
Batrad U, Shelukard S, Kweia G and Storeye G (2004) The role of biopharmaceutics in 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 327 
the development of a clinical nanoparticle formulation of MK-0869: A Beagle dog model 
predicts improved bioavailability and diminished food effect on absorption in human. 
Int} Pharm 285(1-2):135-146. 
40. Liversidge GGCP (1995) Drug particle size reduction for decreasing gastric irritancy and 
enhancing absorption of naproxen in rats. Int J Pharm 125:309-313. 
41. Miiller RH (1997) Preparation in Form of a Matrix Material-Auxiliary Agent Compound Containing 
Optionally an Active Substance: Europe. 
42. Keck CM et al. (2004) Production and Optimisation of Oral Cyclosporine Nanocrystals, in 
AAPS. Baltimore. 
43. Krause K (2004) Herstellung hochfeiner Polymer- und Arzneistoffdispersionen und deren 
Spruhtrocknung, in PhD Thesis Pharmaceutical Technology, Freie Universitat: Berlin. 
44. Khar A (2002) Nanoedge Technologies. Baxter Company Booklet. 

45. Yamaguchi H and Hachioji I (2001) New antifungal agents currently under clinical development. 
Nippon Kagaku Ryoho Gakkai Zasshi 9(49):535-545. 
46. Slain DR, Cleary PD and Chapman SW (2001) Intravenous itraconazole. Annals of pharmacotherapy 
35(6):720-729. 
47. Smith A and Hunneyball LM (1986) Evaluation of poly(lactic acid) as a biodegradable 
drug delivery system for parenteral administration. Int} Pharm 30(2-3):215-220. 
48. Hernandez-Trejo N, Kayser O, Miiller RH and Steckel H (2004) Physical Stability ofBuparvaquone 
Nanosuspensions Following Nebulization with Jet and Ultrasonic Nebulizers. Proceedings 
of the International Meeting on Pharmaceutics, Biopharmaceutics and Pharmaceutical 
Technology, Nuremberg Germany. 
49. Hernandez-Trejo N, Kayser O, Miiller RH and Steckel H (2004) Characterization of nebulized 
buparvaquone nanosuspensions — Effect of nebulization technology. Pharm Res. 
submitted. 
50. Patravale VBD, Abhijit A and Kulkarni RM (2004) Nanosuspensions: A promising drug 
delivery strategy. / Phar Pharmacol 56(7):827-840. 
51. Pignatello R and Puglisi G (2002) Ocular tolerability of Eudragit RS100 and RL100 
nanosuspensions as carriers for ophthalmic controlled drug delivery. / Pharm Sci 
91(12):2636-41. 
52. Bucolo CM, Puglisi G and Pignatello R (2002) Enhanced ocular anti-inflammatory activity 
of ibuprofen carried by an Eudragit RSI 00 nanoparticle suspension. Ophfhal Res 
34(5):319-323. 
53. Miinster UN, Haberland C, Jores A, Mehnert W, Rummel S, Schaller K, Korting M, 
Zouboulis Ch, Blume-Peytavi C and Schafer-Korting M (2005) RU 58841-myristateprodrug 
development for topical treatment of acne and androgenetic alopecia. Die Pharmazie 
60(1):8-12. 
54. Maia C, Mehnert W and Schafer-Korting M (2000) Solid lipid nanoparticles as drug 
carriers for topical glucocorticoids. Int} Pharm 196:165-167. 
55. Miiller RH (2001) Dispersions for the Formulation of Slightly or Poorly Soluble Drugs, in 
PCT/EP01/08726. PharmasSol GmbH Berlin. 
56. Miiller RH et al. (2004) SolEmuls — A novel technology for the formulation of i.v. emulsions 
with poorly soluble drugs. Int} Pharm 269:293-302. 
328 Miiller & Junghanns 
57. Muller RH et al. (2004) SolEmuls-novel technology for the formulation of i.v. emulsions 
with poorly soluble drugs. Int J Pharm 269(2):293-302. 
58. Buttle I (2004) O/W-Emulsionen fiir die intravenose Applikation von Arzneistoffen, in PhD 
Thesis Pharmaceutical Technology, Freie Universitat: Berlin. 
59. Akkar A and Muller RH (2004) Solubilisation by Emulsification. Pharm. Ind. 66(12): 
1537-1544. 
60. Akkar A and Muller RH (2003) Intravenous itraconazole emulsions produced by 
SolEmuls technology. Eur J Pharm Biopharm 56(l):29-36. 
61. Akkar A et al. (2004) Solubilising Poorly Soluble Antimycotic Agents by Emulsification via a 
Solvent-Free Process. AAPS PharmSciTechpending, submitted. 
62. Akkar A and Muller RH (2003) Formulation of intravenous Carbamazepine emulsions 
by SolEmuls technology. Eur J Pharm Biopharm 55(3):305-12. 
63. Akkar A (2004) Poorly Soluble Drugs: Formulation by Nanocrystals and SolEmuls Technologies, 
in PhD Thesis Pharmaceutical Technology. Freie Universitat: Berlin. 
64. Hann IM and Prentice HG (2001) Lipid-based amphotericin B: A review of the last 10 
years of use. Int J Antimicrob Agents (17): 161-169. 
65. Lewis R (2003) Antifungal therapy cost analysis (Patterson T. F. and McGinis M. R., ed.). 
www.doctorfungus.org. 
66. Janoff A et al. (1993) Amphotericin b lipidcomplex (ABLC (TM)): A molecular rationale 
for the attenuation of amphotericin B related toxicities. / Liposome Res 3:451-471. 
67. Davis SS and Washington C (1988) EP 0 296 845 Al. 
15 
Cells and Cell Ghosts as Drug Carriers 
Jose M. Lanao and M. Luisa Sayalero 
1. Introduction 
Microparticle and nanoparticle polymeric systems currently occupy an important 
place in the field of drug delivery and targeting.1 Nevertheless, there are biological 
drug carriers that offer an efficient alternative to such systems. Within the different 
systems of biological carriers, of great importance are cells and cell ghosts, which are 
both efficient and highly compatible systems from the biological point of view, capable 
of providing the sustained release and specific delivery to tissues, organs and 
cells of drugs, enzymatic systems and genetic material. Cell systems such as bacterial 
ghosts, erythrocyte ghosts, polymorphonuclear leukocytes, apoptotic cells, 
tumor cells, dendritic cells, and more recently, genetically engineered stem cells, 
are all examples of how cell systems of very diverse nature can be suitably manipulated 
and loaded with drugs and other substances, to permit specific drug delivery 
in vivo with important therapeutic applications.2-8 Cell carriers for drug delivery are 
used in very different applications such as cancer therapy, cardiovascular disease, 
Parkinson's, AIDS, gene therapy, etc. Table 1 shows the classification of biological 
carriers for drug delivery based on the use of cells and cell ghosts. 
2. Bacterial Ghosts 
Bacterial ghosts are intact, non-living, non-denatured bacterial cell envelopes 
devoid of cytoplasmic contents. They are created by lysis of bacteria, but maintain 
329 
330 Lanao & Sayalero 
Table 1 Kinds of cells and cell ghosts used for drug and gene delivery. 
Cell carrier 
Bacterial ghost 
Erythrocyte ghost 
Engineered stem cells 
Polymorphonuclear 
leucocytes 
Apoptopic cells 
Tumor cells 
Denditric cells 
Target 
Tissues, macrophages, cells 
RES, macrophages 
Tumor cells, T cells, 
macrophages 
Tissues 
Tumor cells 
Tumor cells 
T cells 
Encapsulated substance 
Drugs, vaccines, genetic material 
Drugs, enzymes, peptides 
Genetic material 
Drugs 
Drugs 
Drugs 
Drugs 
their cellular morphology and native surface antigenic structures, including their 
bioadhesive properties.3,9 
Bacterial ghosts allow the encapsulation of drugs and other substances, and 
their specific attachment to mammalian tissues and cells. This kind of cell carrier acts 
as a true drug delivery system, allowing the permanency of drugs in the systemic 
circulation to be increased together with tissue-specific targeting. They are thus a 
promising alternative to conventional drug delivery systems such as liposomes or 
nanoparticles. 
The main advantages of bacterial ghosts as delivery systems are the fact that 
they are non-living, i.e. they can act as delivery systems of drugs, antigens or DNA; 
allow specific delivery to different tissues and cell types; and are well captured by 
phagocytic cells and antigen-presenting cells as dendritic cells. Among the drawback 
of bacterial ghosts is the possibility that they might revert to being virulent, the 
possibility of horizontal gene transfer, the stability of the recombinant phenotype, 
and pre-existing immunity against the carrier used.10 
Usually, bacterial ghosts are produced by protein E-mediated lysis of Gramnegative 
bacteria.11 The production of bacterial ghosts is based on the controlled 
expression of the bacteriophage PhiX174-derived lysis gene E. Expression of this 
gene from a plasmid in Gram-negative bacteria leads to the formation of a transmembrane 
lysis tunnel structure that penetrates the inner and outer membranes, 
and is formed by protein E with border values fluctuating between 40-200 nm 
in diameter. Protein E is a hydrophobic protein localized exclusively in the cell 
envelope.12 E-mediated lysis has been achieved in many Gram-negative bacteria.13 
Scanning electron micrographs of E-lysed cells reveal that bacterial ghosts contain 
only one E hole in a bacterial ghost, although in a few cases, there are two holes. 
The cytoplasm is expelled as a consequence of the high osmotic pressure inside the 
Cells and Cell Ghosts as Drug Carriers 331 
cell. The collapse of membrane potential and the release of cytoplasmic components 
such as proteins, nucleic acids, etc occur simultaneously.14 In the case of strains of 
E. coli, this effect occurs within a period of 10 min after the induction of expression.15 
The resulting empty bacterial cell envelope is considered a bacterial ghost. Bacterial 
ghosts show all the morphological, structural and immunogenic properties of 
a living cell.9-15-17 Since bacterial ghosts are derived from Gram-negative bacteria 
that are able to adhere to structures such as fimbriae and lipopolysaccharide, they 
are used for specific binding to human tissue.18 
Bacterial ghost drug-loading is accomplished by the suspension of lyophilised 
bacterial ghosts in a buffered medium containing the drug. The ghosts are then 
subjected to an incubation process varying from 5 to 30 min at 24°C. They are then 
washed to remove excess drug.18'19 
In order to prevent rapid leakage of loaded water-soluble drugs or other substances, 
the bacterial ghosts are sealed by fusion of the cell membrane with membrane 
vesicles at the edges of the lysis pore. For the sealing step, the bacterial ghosts 
suspension is incubated in the fusion buffer at 28°C for 10 min.18 Figure 1 shows a 
scheme of the production of bacterial ghosts by protein E-mediated bacterial lysis. 
The in vitro release of drugs from loaded bacterial ghosts is performed from a 
suspension of drug-loaded bacterial ghosts that is dialysed through a membrane 
suitable for excluding the ghosts. Dialysis is performed at 28°C in PBS buffer.19 The 
concentrations of drug released into the buffer at preset times are quantified using 
an appropriate analytical technique. 
In studies addressing the adherence and capture of loaded bacterial ghosts by 
target cells such as macrophages, human colorectal adenocarcinoma cells (Caco-2) 
or dendritic cells, fluorescent markers such as fluorescein isothiocyanate (FITC) 
are used. These allow adherence to be assessed using fluorescence microscopy 
and flow cytometry techniques.18,19 Macrophages internalize bacterial ghosts to 
a greater extent than Caco-2 cells.18,19 Studies carried out using confocal laser scanning 
microscopy with M. haemolytica ghosts loaded with Doxorubicin have shown 
that the drug was associated with the ghosts membranes and the inner lumen.19 
Denditric cells that are professional phagocytic cells displaying the phagocytic 
capacity of antigens also have a good capacity for capturing bacterial ghosts, allowing 
the latter to be used as a vehicle for immunization and immunotherapy.20 
2.1. Application of bacterial ghosts as a delivery system 
Bacterial ghosts have important therapeutic applications. They can be loaded 
with drugs, proteins and other substances, and can be targeted selectively to 
macrophages, tumors or endothelial cells.10,19 
332 Lanao & Sayalero 
Cytoplasmic content 
AAAAAAA 
Inner Membrane 
Outer Membrane 
GRAM-NEGATIVE BACTERIA 
Cytoplasmic Membrane 
Protein E-mediated lysis 
E hole (40-200 nm) 
DRUG-LOADING BACTERIAL GHOST 
DRUG-LOADED BACTERIAL GHOST 
Fig. 1. Production and drug loading of bacterial ghosts. 
Bacterial ghosts have been used as efficient drug delivery systems10 in the 
field of anti-cancer drugs.18 Bacterial ghosts obtained have been used as a delivery 
system of doxorubicin to human colorectal carcinoma cells. Cytotoxicity assays 
revealed that doxorubicin-loaded ghosts show better antiproliferative capacity in 
Caco-2 cells than when free doxorubicin is used at the same concentration.18 Experiments 
have also been carried out using E.coli ghosts containing streptavidin, in order 
to increase the affinity of streptavidin for biotinylated compounds. Streptavidinloaded 
ghosts permit specific targeting to mucosal surfaces of the gastrointestinal 
and respiratory tracts, and also to phagocytic cells.3 Bacterial ghosts have been used 
as veterinary vaccines for the immunization of different animal species.9 
Pasteurella multocida is a pathogen that causes morbidity and mortality in rabbits 
and its importance as a human pathogen has also been recognized. P. multocida 
Cells and Cell Ghosts as Drug Carriers 333 
ghosts have been used to immunize rabbits and mice.17 Similar results have been 
obtained in the immunization of cattle against pasteurellosis using Pasteurella 
haemolytica ghosts.11 
Actinobacillus pleuropneumoniae is a highly contagious microorganism and is 
the cause of porcine pleuropneumonia, infecting 30-50% of pig populations. However, 
Actinobacillus pleuropneumoniae vaccines provide limited protection, since they 
decrease mortality but not morbidity in swine. Comparative studies have been carried 
out on immunization using a aerosol infection model for pigs vaccinated with 
loaded-ghosts or formalin inactivated Actinobacillus pleuropneumoniae bacterins. The 
results obtained showed that immunization with bacterial ghosts is more efficient 
in protecting pigs than bacteria.21,22 
Bacterial ghosts can also be used as carriers of therapeutic DNA or RNA.3/13 
The use of nucleic acid vaccines currently offers a technique for the development of 
prophylactic or therapeutic vaccines, based on the use of DNA plasmids to induce 
immune responses by direct administration of DNA-encoding antigenic proteins 
into animals, and this is also suitable for the induction of cytotoxic T cells.23,24 
Bacterial ghosts loaded with DNA produce a high level of gene expression. They 
can be used to enhance the mucosal immune response to target antigens expressed 
in the bacterial ghost system. They can also be used for the specific targeting of 
DNA-encoded antibodies to primary antigens located in cells.13 Ghosts of Vibrium 
cholerae have been tested as antigen carriers of Chlamidia trachomatis as potential 
vaccines for the control of genital infections produced by this bacteria. Recombinant 
Vibrium cholerae ghosts, previously cloned with a major outer membrane protein of 
C. trachomatis, afforded a high level of protective immunity against Chlamydia in a 
murine model.25,26 Mannheimia haemolytica is a pathogen that causes ovine mastitis. 
M. haemolytica ghosts loaded with plasmid DNA stimulate the elicitation of efficient 
immune responses in mice, with no symptoms of acute or subacute toxicity during 
the experiment.27 
3. Erythrocyte Ghosts 
Erythrocytes constitute the largest population of blood cells and are produced in 
the bone marrow. They are mature blood cells that produce haemoglobin and carry 
out the exchange of oxygen and carbon dioxide between the lungs and the body 
tissues. 
The term "erythrocyte ghost" attempts to define the resulting cell-like structure 
when erythrocytes are subjected to a reversible process of osmotic lysis.28 For more 
than 30 years, many studies, both in vivo and in vitro, have been carried out to explore 
the use of erythrocyte ghosts as delivery systems of drugs and other substances.2 
334 Lanao & Sayalero 
Erythrocyte ghosts are obtained from fresh erythrocytes coming from human 
blood or the blood of different animal species such as the rat, mouse, rabbit, etc, and 
are loaded with different types of substance, mainly drugs, peptides and enzymes, 
using different encapsulation methods. The most frequent methods for collecting 
erythrocyte ghosts are osmosis-based methods such as hypotonic dialysis.2,29 
Autologous erythrocyte ghosts offer a drug delivery system that can act as a 
reservoir of the drug or substance encapsulated, providing the sustained release 
of the drug into the organism together with selective targeting of the drugs to the 
reticuloendothelial system (RES) of the liver, spleen and bone marrow.2 
The main advantages of carrier erythrocytes as drug delivery systems are their 
high degree of biocompatibility, the possibility of encapsulating the drug in a small 
amount of cells, the sustained release of the encapsulated drug or substance into 
the body, the selective targeting to the RES, and the possibility of encapsulating 
substances of high molecular weight such as peptides. Among the drawbacks of 
these systems are the rapid leakage of some drugs out of the loaded erythrocytes 
and other problems related to their standardized preparation, storage and potential 
contamination.2 
Erythrocyte ghosts can be obtained by diverse procedures such as hypotonic 
dilution, hypotonic pre-swelling, osmotic pulse, hypotonic hemolysis, hypotonic 
dialysis, electroporation, drug-induced endocytosis and chemical methods.2,30 Of 
the different ways of obtaining carrier erythrocytes, hypotonic dialysis is undoubtedly 
the most frequently used encapsulation method. The reasons why it is so popular 
are its simplicity, its ease of application for a large number of drugs, enzymes 
and other substances, and because it is the method that best conserves the morphological 
and haematological properties of the erythrocyte ghosts obtained. 
Hypotonic dialysis is based on the exposure of red cells to the action of a hypotonic 
buffer, inducing cell swelling and the formation of pores that permit the drug 
to enter erythrocytes by means of a passive mechanism. Figure 2 shows a scheme 
of the production of erythrocyte ghosts using a hypotonic dialysis method. 
Morphological inspection of erythrocyte ghosts is usually performed using 
transmission (TEM) or scanning (SEM) electron microscopy.2 Some physical parameters 
of red cell membranes can also be studied from the diffusion of haemoglobin.28 
The haemolytic methods employed in the production of erythrocyte ghosts normally 
affect the haemolytic volume, surface area and tension.28 Figure 3 shows the 
morphological changes observed by SEM that occur in amikacin-loaded erythrocytes 
due to hypotonic dialysis.31 
Haematological parameters determine the effects of the procedure used to collect 
erythrocyte ghosts on their haematological properties. Among others, parameters 
such as reduced glutathione (GSH), mean corpuscular volume (MCV) or red 
cell distribution width (RDW), may be evaluated using a haematology analyzer. 
Cells and Cell Ghosts as Drug Carriers 335 
Dialysis bag 
SG 
ERYTHROCYTES 
Erythrocytes 
suspension T \ 
Drug 
ANNEALING 
(Isotonic buffer) 
RESEALFNG 
(Hypertonic buffer) 
(10 min, 37°C, pH 7.4) (30 min, 37°C, pH 7.4) 
Hypotonic buffer 
HYPOTONIC DIALYSIS 
(45 min, 4"C, pH 7.4) 
DRUG-LOADED GHOST 
ERYTHROCYTES 
Fig. 2. Production and drug loading of erythrocyte ghosts using a hypotonic dialysis 
method. 
o V 
CONTROL 
ERYTHROCYTES 
V'^C 
AMIKACIN LOADED 
ERYTHROCYTES 
Fig. 3. SEM micrographs of amikacin carrier erythrocytes obtained by hypotonic dialysis31 
(Copyright 2005 from Encapsulation and in vitro Evaluation of Amikacin-Loaded Erythrocytes 
by C. Gutierrez Millan. Reproduced by permission of Taylor & Francis Group, LLC, 
http: / / www. taylorandfrancis .com). 
Erythrocyte ghosts obtained by hypotonic dialysis show a decrease in the mean corpuscular 
volume and an increase in size dispersion.28'29 Erythrocyte ghosts show a 
greater degree of haemolysis than normal erythrocytes.29 
3.1. Applications of erythrocyte ghosts as a delivery system 
Erythrocyte ghosts can be used as potential drug delivery systems for enzymes, 
proteins and peptides, allowing sustained release into the systemic circulation and 
the delivery of these substances into the RES.2 
In vitro release of drugs from loaded erythrocyte ghosts is usually tested using 
autologous plasma or an isoosmotic buffer at 37°C; alternatively, a dialysis bag 
may be used.32 The in vitro release of drugs and substances from loaded erythrocytes 
is usually a first-order process, suggesting that the drug crosses the plasma 
membrane through a passive diffusion mechanism.33 However, zero-order release 
336 Lanao & Sayalero 
kinetics from loaded erythrocytes has also been described.34 In vitro studies about 
the release kinetics of different drugs, enzymes and peptides from loaded erythrocytes 
have shown a slow release of the encapsulated substance.2 
When loaded erythrocyte ghosts are administered in vivo, changes in the pharmacokinetics 
of the encapsulated drugs occur, involving a systemic drug clearance 
related to the biological half-life of the erythrocytes.35 Increased serum half-lives 
and the areas under the curve of drugs encapsulated in loaded erythrocyte ghosts, 
in comparison with the free drug, have been observed in animals and humans.36,37 
At the same time, erythrocyte ghosts show a greater accumulation in tissues such 
as liver and spleen.38,39 
Surface treatment of erythrocyte ghosts with substances such as glutaraldehyde, 
ascorbate, Fe(+2), diamide, band 3-cross-linking reagents, trypsin, phenylhydrazine 
and the N-hydroxysuccinimide ester of biotin (NHS-biotin), enhances 
the recognition of erythrocyte ghosts by macrophages in vitro and liver targeting 
in vivo.40~i3 
Red cells may be used as carriers for some drugs such as antineoplastics, antiinfective 
agents, antihypertensives, corticosteroids, etc.2 Thus, carrier erythrocytes 
have been widely studied as delivery systems of antineoplastic drugs for targeting 
the RES located in organs such as liver and spleen. 
Different antineoplastic drugs have been encapsulated in erythrocyte ghosts 
in both in vitro and in vivo experiments.2 Increases have been obtained in average 
survival times in the treatment of mice bearing hepatomas, using methotrexateloaded 
carrier erythrocytes.44 Better recognition and capture of erythrocyte ghosts 
by macrophages have been obtained by using biotinylated erythrocytes containing 
methotrexate,45 by alterations to the membrane using band-3 cross-linkers of erythrocyte 
ghosts containing etoposide,46 or by treatment of erythrocytes containing 
doxorubicin with glutaraldehyde.47 
Anti-infective agents such as gentamicin, metronidazole, primaquine or imizol 
have also been encapsulated in erythrocytes.2 Human erythrocytes containing 
gentamicin have proven to act as an efficient slow-release system in ofco.48,49 
Erythrocyte ghosts containing dexamethasone have been used in vivo in rabbits 
and humans. A sustained release of dexamethasone in vivo in animals and humans 
was observed using carrier erythrocytes. An increased anti-inflammatory effect of 
the drug using carrier erythrocytes was observed in rabbits.50,51 
Moreover, new prodrugs of anti-opioid drugs such as naltrexone and naloxone 
have been encapsulated in erythrocytes to solve stability problems of the primary 
drug within the erythrocyte. The encapsulated prodrugs are transformed into the 
active compound, following their release from erythrocyte ghosts.52 
In the fields of biochemistry and enzymatic therapeutics, the encapsulation 
of enzymes in erythrocytes has been studied in some depth. Enzymatic 
Cells and Cell Ghosts as Drug Carriers 337 
deficiencies or the treatment of specific illnesses may be approached using carrier 
erythrocytes loaded with enzymes. The encapsulation of enzymes in erythrocytes 
solves some of the problems associated with enzyme therapy, such as 
the short half-life deriving from the action of plasma proteases, intolerant reactions, 
and the immunological disorders or allergic problems associated with 
the use of enzymes in therapeutics. In vitro or in vivo studies with enzyme 
carrier erythrocytes have been performed using L-asparaginase,53 hexokinase,54 
alcohol dehydrogenase,55 aldehyde dehydrogenase,56 alcohol oxidase,57 glutamate 
dehydrogenase,58 uricase,59 urokinase,60 lactate 2-mono oxigenase,61 
arginase,62 rhodanase,63 recombinant phosphotriestearase,64 delta-aminolevulinate 
dehydratase,65 urease,66 pegademase,67 brinase68 and alglucerase.69 One of the best 
examples of the use in therapeutics of carrier erythrocytes containing enzymes, is 
that of L-asparaginase encapsulated in human erythrocytes. This has been successfully 
used in the treatment of acute lymphoblastic leukaemia in paediatrics.70 
Erythrocyte ghosts may act as carrier systems for the delivery of peptides 
and proteins. One of the main therapeutic applications of carrier erythrocytes in 
this field is that of anti-HIV peptides. Nucleoside analogues successfully inhibit 
the replication of immunodeficiency virases. In view of the importance of the 
monocyte-macrophage system in infection by HIV-1, it would be of maximum 
therapeutic interest to have available, the specific delivery of these therapeutic 
peptides into macrophages, which act as an important reservoir for the virus. Carrier 
erythrocytes containing anti-HIV peptides such as azidothimidine (AZT) and 
didanosine (DDI), significantly reduced the pro-viral DNA content in comparison 
with the administration of free peptides in a murine AIDS model.71 Similar 
results have been obtained with 2',3'-dideoxycytidine 5'-triphosphate'(ddCTP),72 
2',3'-dideoxycytidine (ddCyd)73 and AZT prodrugs74 encapsulated in erythrocytes. 
Anti-neoplastic peptides such as 2-fluoro-ara-AMP (fludarabine) and 5'- 
fluoro-2'-deoxyuridine 5'-monophosphate (FdUMP), a pro-drug of 5-fluro-2'- 
deoxyuridine (FdUrd), have been encapsulated in human carrier erythrocytes, 
behaving as a slow-release delivery system.75,76 
Macrophage uptake in vitro of antisense oligonucleotides may be increased by 
using carrier erythrocytes.77,78 Other peptides, such as erythropoietin,79 heparin,80 
dermaseptin S3,81 interleukin-382 or vaccines,83 have also been encapsulated in erythrocytes 
to increase their stability,84 acting as a slow release system with a prolonged 
half-life,80,84 or for their specific targeting to bacterial membranes.85 
Erythrocyte ghost derivatives can also be used as drug delivery systems. 
Nanoerythrosomes are erythrocyte membrane derivatives formed by spheroid vesicles, 
obtained by consecutive extrusion under nitrogen pressure through a polycarbonate 
filter membrane of a erythrocyte ghost suspension to produce small vesicles 
having an average diameter of 100 nm. In vitro and in vivo studies, carried out with 
338 Lanao & Sayalero 
nanoerythrosomes loaded with daunorubicin, have shown that when linked covalently 
to nanoerythrosomes, the drug produces slow release of daunorubicin to the 
organism over a prolonged period of time and also that, in comparison with the free 
drug, cytotoxicity is greater.86 The advantage of nanoerythrosomes, as compared 
with erythrocyte ghosts as drug delivery system, is that the former are able to escape 
from the reticuloendothelial system faster.86,87 In vitro studies have shown that the 
nanoerythrosome-daunorubicin complex is rapidly adsorbed and phagocytosed by 
macrophages.88 Liver, spleen and lungs are the organs in which nanoerythrosomes 
show the greatest capacity of accumulation.89 
Another derivative of erythrocyte ghosts are reverse biomembrane vesicles 
loaded with drugs.90 Reverse biomembrane vesicles are produced by spontaneous 
vesiculation of the ghost erythrocyte membrane by endocytosis, using an appropriate 
vesiculating medium, producing small vesicles containing the drug within the 
parent ghost. In vivo studies carried out using reverse biomembrane vesicles from 
erythrocyte ghosts loaded with doxorubicin in rats have revealed increases in the 
half-life and bioavailability of the drug, the liver and spleen, being the main organs 
for the clearance of this drug delivery system.90 
4. Stem Cells 
In gene therapy, a therapeutic transgene is introduced into the patient with a view of 
supplementing the functions of an abnormal gene. To achieve the delivery of genetic 
material into the target cell, it is necessary to have a suitable carrier. One important 
aim in the field of gene therapy is the design and development of gene carriers that 
encapsulate and protect the nucleic acid, and selectively release the vector/nucleic 
acid complex to the target tissue, so that the genetic material will be released at the 
cellular level later. In practice, there are several ways to achieve this. The first is 
through the use of modified viruses containing the genetic material of interest. The 
use of viruses for gene delivery has some drawbacks since it is limited to specific 
cells susceptible to being infected by the virus, and also the administration itself 
of the virus, has some immunological problems among others.91-93 The second 
alternative is to use living cells modified genetically, such as stem cells, to deliver 
transgenic material into the body.8,94 
Stem cell therapy is a new form of treatment, in which cells that have died or 
lost their function are replaced by healthy adult stem cells. One advantage of this 
kind of cell is that it is possible to use samples from adult tissues or cells from the 
actual patient, for culture and subsequent implantation. 
Within the framework of stem cell research, the use of stem cells as delivery 
systems is a novel and attractive technique in the field of gene therapy, in which the 
cells of the patients themselves are genetically engineered, in order to introduce a 
therapeutic transgene used to deliver the genetic material. A promising therapeutic 
Cells and Cell Ghosts as Drug Carriers 339 
strategy is the use of stem cells such as lymphocytes or fibroblasts as drug delivery 
systems. Experimental studies using stem cells as such systems have been tested 
in different therapeutic applications, especially in the field of cancer therapy. Considering 
the affinity of stem cells for tumor tissue, engineered stem cells have been 
successfully used for direct drug delivery to cancer cells.8'94 In vitro cultures have 
been made of human mesenchymal stem cells from bone marrow that are transduced 
with an adenovirus vector carrying the human interferon beta-gene, which 
exerts therapeutic action against cancer. Engineered stem cells administered in vivo 
allow the delivery of the genetic material to cancer cells. This new drug delivery 
system has proven to be efficient in the treatment of experimental neoplasms, such 
as lung cancer, in mice.94 Figure 4 shows a scheme of the application of stem cells 
as carriers for gene delivery in experimental cancer therapy. 
In vivo studies have also been carried out with neural stem cells engineered 
using adenoviral vectors to express interleukm-12, an oncolytic gene, whose efficiency 
has been demonstrated in the treatment of intracranial malignant gliomas 
in mice.8,95 
The used of haematopoietic stem cells has allowed antiviral genes to be introduced 
in both T cells and macrophages for the treatment of AIDS.96 The use of 
stem cells as vehicles for gene therapy has also been suggested for the treatment of 
ischaemic heart disease,97 
Stem cells have also been employed in the field of antiepileptic therapy. Glial 
precursor cells, which release adenosine, have been derived from adenosine kinase 
embryonic stem cells. In these experiments, the fibroblasts were engineered to 
release adenosine by inactivating adenosine metabolising enzymes. After encapsulation 
within polyethersulfone hollow-fibre capsules, and the introduction into 
Mesenchymal 
Stem cells Interferon 
Beta 
Recombinant 
adenovirus 
Engineered i„ vitro 
Stem cells expansion 
'. 
Ficoll ; ', 
Bone 
marrow 
Fig. 4. Application of stem cells as carriers for gene delivery in experimental cancer. 
340 Lanao & Sayalero 
the brain ventricles in a rat epilepsy model, the local release of adenosine allows 
drug-resistant focal epilepsy to be treated. These engineered cells were shown to 
suppress seizure activity.98-99 
5. Polymorphonuclear Leucocytes 
Polymorphonuclear leucocytes (PMN) can be used as carriers of antibiotics in view 
of their selective targeting to sites of infection. Simply incubating PMN in the presence 
of high concentrations of antibiotic for 1 hr at 37° C guarantees cell loading with 
the antibiotic. PMN loaded with the macrolide azithromycin have been found to be 
efficient in an in vitro model that permits the delivery of the antibiotic in a bioactive 
form to Chlamydia inclusions in polarized human endometrial epithelial (HEC-1B) 
cells infected with Chlamydia trachomatis. PMN carriers allow the accumulation of 
large amounts of antibiotic in endometrial epithelial cells and its retention over 
long periods of time.4 
6. Apoptopic Cells 
Programmed cellular death or apoptosis is a process that is controlled genetically 
in which the cells induce their own death in response to different types of stimulus 
such as the binding of death-inducing ligands to cell surface receptors. 
A new strategy for drug delivery, called apoptopic induced drug delivery 
(AIDD), allows drug delivery to tumor cells upon the initiation of apoptosis by 
using a biological mechanism to achieve drug delivery.5 This new system is based 
on the fact that apoptosis produces many changes in cell morphology that can be 
taken advantage of to achieve drug delivery. 
Apoptosis is reflected in enhanced membrane permeability, which favors the 
release of the encapsulated drug from the apoptotic cells to the tissue. Phagocytosis 
of drug- loaded apoptotic carrier cells by tumor cells facilitates the localization 
of the drug within the tumor cell. One advantage of the apoptotic induced drug 
delivery system (AIDD) is that the drug carrier cells may be genetically engineered 
to modify their properties. 
In vitro studies have been performed using S49 mouse lymphoma cells in which 
apoptosis is produced by exposure to dexamethasone. The cytotoxicity of RG-2 cells 
caused by temazolamide-loaded-S49 apoptotic cells was from 4 to 7 times higher 
than that of control temazolamide-loaded S49 cells.5 
7. Tumor Cells 
A novel strategy for drug delivery based on the use of cell systems is the drugloaded 
tumor cell system (DLTC), developed for drug delivery and targeting in 
Cells and Cell Ghosts as Drug Carriers 341 
lung metastasis.6'100 The tumor cells as drug carriers permit drug targeting to the 
blood-borne cancerous cells and the lungs as potential metastatic organs. In practice, 
there is affinity between the plasma membrane of malignant tumor cells and the 
metastatic addressins expressed by the endothelial cells of the targeted organ.6101 
In vivo studies have been carried out with DLTC based on Doxorubicin-loaded 
B16-F10 murine melanoma cells. Doxorubicin accumulation in the mouse lung was 
several times higher than that seen after administering free Doxorubicin.6 
8. Dendritic Cells 
Dendritic cells (DC) are antigen-presenting cells. They ingest antigen by phagocytosis, 
degrade it, and present fragments of the antigen at their surface. Dendritic cells 
have huge potential for immunization against a broad variety of diseases, because 
they travel throughout the body in search of pathogens indicative of infection or 
disease. They are very important for the induction of T cell responses, which result 
in cell-mediated immunity. 
Selective targeting of drugs incorporated in dendritic cells to T cells allows 
the response of these cells to be manipulated in vivo. It has been shown that when 
incorporated into dendritic cells, the drug O-galactosylceramide improves their 
anti-tumor activity.7 
9. Conclusions 
This chapter has focused on the use of cells and cell ghosts as delivery systems of 
drugs, enzymes or therapeutic genes. The use of carrier cells such as bacterial ghosts, 
erythrocyte ghosts and engineered stem cells, for drug delivery and targeting are 
reviewed among others. Their high biocompatibility, together with their capacity 
for selective delivery and targeting in cells and specific tissues mean that these types 
of carrier are promising alternatives to the use of nano- and microparticle systems, 
with applications in the fields of interest such as cancer therapy, cardiovascular 
therapy, AIDS, gene therapy, etc. As an alternative to the use of cell carriers, modified 
viruses can also be used as drug delivery systems, especially in the field of gene 
therapy. Despite their potential interest, clinical studies with these types of carrier 
are still very limited, although in the near future, increase in the use and therapeutic 
applications of cell delivery systems is expected. 
Acknowledgments 
This chapter was supported in part by a project of the National Research and Development 
Plan (Project: SAF 2001-0740). 
342 Lanao & Sayalero 
References 
1. Mainardes RM and Silva LP (2004) Drug delivery systems: Past, present, and future. 
Curr Drug Targets 5:449-455. 
2. Gutierrez-Millan C, Sayalero ML, Castaneda AZ and Lanao JM (2004) Drug, enzyme 
and peptide delivery using erythrocytes as carriers. / Control Rel 95:27-49. 
3. Huter V, Szostak MP, Gampfer J, Prethaler S, Wanner G, Gabor F and Lubitz W 
(1999) Bacterial ghosts as drug carrier and targeting vehicles. / Control Rel 61: 
51-63. 
4. Paul TR, Knight ST, Raulston JE and Wyrick PB (1997) Delivery of azithromycin to 
Chlamydia trachomatis-infected polarized human endometrial epithelial cells by polymorphonuclear 
leucocytes. / Antimicrob Chemother 39:623-630. 
5. Man J and Gallo JM (1998) Delivery of cytotoxic drugs from carrier cells to tumour cells 
by apoptosis. Apoptosis 3:195-202. 
6. Shao J, DeHaven J, Lamm D, Weissman DN, Runyan K, Malanga CJ, Rojanasakul Y 
and Ma JK (2001) A cell-based drug delivery system for lung targeting: II. Therapeutic 
activities on B16-F10 melanoma in mouse lungs. Drug Del 8:61-69. 
7. Fujii S, Shimizu K, Kronenberg M and Steinman RM (2002) Prolonged IFN-gammaproducing 
NKT response induced with alpha-galactosylceramide-loaded DCs. Nat 
Immunol 3:867-874. 
8. Yang SY, Liu H and Zhang JM (2004) Gene therapy of rat malignant gliomas using 
neural stem cells expressing IL-12. Cell Biol 23:381-389. 
9. Jalava K, Hensel A, Szostak M, Resch S and Lubitz W (2002) Bacterial ghosts as vaccine 
candidates for veterinary applications. / Control Rel 85:17-25. 
10. Tabrizi CA, Walcher P, Mayr UB, Stiedl T, Binder M, McGrath J and Lubitz W (2004) 
Bacterial ghosts-biological particles as delivery systems for antigens, nucleic acids and 
drugs. Curr Opin Biotechnol 15:530-537. 
11. Marchart J, Dropmann G, Lechleitner S, Schlapp T, Wanner G, Szostak MP and Lubitz W 
(2003) Pasteurella multocida- and Pasteurella haemolytica-ghosts: New vaccine candidates. 
Vaccine 21:3988-3997. 
12. Altman E, Young KD, Garrett J, Altman R and Young R (1985) Subcellular localization 
of lethal lysis proteins of bacteriophages X and X174. / Virol 53:1008-1011. 
13. Jalava K, Eko FO, Riedmann E and Lubitz W (2003) Bacterial ghosts as carrier and 
targeting systems for mucosal antigen delivery. Exp Rev Vaccines 2:45-51. 
14. Witte A and Lubitz W (1989) Dynamics of PhiX174 protein E-mediated lysis of 
Escherichia coli. Eur } Biochem 393. 
15. Witte A, Wanner G, Blasi U, Halfmann G, Szostak M and Lubitz W (1990) Endogenous 
transmembrane tunnel formation mediated by phi XI74 lysis protein E. / Bacteriol 
172:4109-4114. 
16. Witte A, Wanner G, Sulzner M and Lubitz W (1992) Dynamics of PhiX174 protein Emediated 
lysis of Escherichia coli. Protective immunity against pasteurellosis in cattle, 
induced by Pasteurella haemolytica ghosts. Arch Microbiol 381. 
Cells and Cell Ghosts as Drug Carriers 343 
17. Marchart J, Rehagen M, Dropmann G, Szostak MP, Alldinger S, Lechleitner 
S, Schlapp T, Resch S and Lubitz W (2003) Protective immunity against pasteurellosis 
in cattle, induced by Pasteurella haemolytica ghosts. Vaccine 21:1415- 
1422. 
18. Paukner S, Kohl G, Jalava K and Lubitz W (2003) Sealed bacterial ghosts-novel targeting 
vehicles for advanced drug delivery of water-soluble substances. / Drug Targ 
11:151-161. 
19. Paukner S, Kohl G and Lubitz W (2004) Bacterial ghosts as novel advanced drug delivery 
systems: Antiproliferative activity of loaded doxorubicin in human Caco-2 cells. 
/ Control Rel 94:63-74. 
20. Halsberger AG, Khol G, Felnerova D, Mayr UB, Furst-Ladani S and Lubitz W (2000) 
Activation, stimulation and uptake of bacterial ghosts in antigen presenting cells. 
/ Biotechnol 83:57-66. 
21. Huter V, Hensel A, Brand E and Lubitz W (2000) Improved protection against lung colonization 
by Actinobacillus pleuropneumoniae ghosts: Characterization of a genetically 
inactivated vaccine. / Biotechnol 83:161-172. 
22. Hensel A, Huter V, Katinger A, Raza P, Strnistschie C, Roesler U, Brand E and Lubitz W 
(2000) Intramuscular immunization with genetically inactivated (ghosts) Actinobacillus 
pleuropneumoniae serotype 9 protects pigs against homologous aerosol challenge 
and prevents carrier state. Vaccine 18:2945-2955. 
23. Donnelly JJ, Ulmer JB, Shiver JW and Liu MA (1997) DNA vaccines. Annu Rev Immunol 
15:617-648. 
24. Felnerova D, Kudela P, Bizik J, Haslberger A, Hensel A, SaalmuUer A and Lubitz W 
(2004) T cell-specific immune response induced by bacterial ghosts. Med Sci Monit 
10:BR362-370. 
25. Eko FO, Lubitz W, McMillan L, Ramey K, Moore TT, Ananaba GA, Lyn D, Black CM 
and Igietseme JU (2003) Recombinant Vibrio cholerae ghosts as a delivery vehicle for 
vaccinating against Chlamydia trachomatis. Vaccine 21:1694-1703. 
26. Eko FO, He Q, Brown T, McMillan L, Ifere GO, Ananaba GA, Lyn D, Lubitz W, Kellar KL, 
Black CM and Igietseme JU (2004) A novel recombinant multisubunit vaccine against 
Chlamydia. ] Immunol 173:3375-3382. 
27. Ebensen T, Paukner S, Link C, Kudela P, de Domenico C, Lubitz W and Guzman CA 
(2004) Bacterial ghosts are an efficient delivery system for DNA vaccines. / Immunol 
172:6858-6865. 
28. Hoffman J (1992) Magnani M and DeLoach JR (eds.) The Use of Reseated Erythrocytes 
as Carriers and Bioreactors. Advances in Experimental Medicine and Biology 326, Plenum 
Press: New York, pp. 1. 
29. Gutierrez Millan C, Zarzuelo Castaneda A, Sayalero Marinero ML and Lanao JM (2004) 
Factors associated with the performance of carrier erythrocytes obtained by hypotonic 
dialysis. Blood Cells Mol Dis 33:132-140. 
30. Hamidi M and Tajerzadeh H (2003) Carrier erythrocytes: An overview. Drug Deliv 
10:9-20. 
344 Lanao & Sayalero 
31. Gutierrez-Millan C, Arevalo M, Zarzuelo A, Gonzalez F> Sayalero ML and Lanao JM 
(2005) Encapsulation and in vitro evaluation of amikacin-loaded erythrocytes. Drug Del 
12:409^16. 
32. Pitt E, Johnson CM and Lewis DA (1983) Encapsulation of drugs in intact erythrocytes: 
An intravenous delivery system. Biochem Pharmacol 32:3359-3368. 
33. Ogiso T, Iwaki M and Ohtori A (1985) Encapsulation of dexamethasone in rabbit erythrocytes, 
the disposition in circulation and anti-inflammatory effect. / Pharmacobio- 
Dyn 8:1032-1040. 
34. Hamidi M, Tajerzadeh H, Dehpour AR, Rouini MR and Ejtemaee-Mehr S (2001) In 
vitro characterization of human intact erythrocytes loaded by enalaprilat. Drug Del 8: 
223-230. 
35. Bax BE, Bain MD, Talbot PJ, Parker-Williams EJ and Chalmers RA (1999) Survival of 
human carrier erythrocytes in vivo. Clin Sci 96:171-178. 
36. Magnani M, Rossi L, Fraternale A, Bianchi M, Antonelli A, Crinelli R and Chiarantini L 
(2002) Erythrocyte-mediated delivery of drugs, peptides and modified oligonucleotides. 
Gene Titer 9:749-751. 
37. Tonetti M, Astroff AB, Satterfield W, De Flora A, Benatti U and DeLoach JR (1991) 
DeLoach, Pharmacokinetic properties of doxorubicin encapsulated in glutaraldehydetreated 
canine erythrocytes. Am } Vet Res 52:1630-1635. 
38. Lizano C, Perez MT and Pinilla M (2001) Mouse erythrocytes as carriers for coencapsulated 
alcohol and aldehyde dehydrogenase obtained by electroporation in vivo survival 
rate in circulation, organ distribution and ethanol degradation. Life Sci 68:2001-2016. 
39. Alvarez FJ, Herraez A, Murciano JC, Jordan JA, Diez JC and Tejedor MC (1996) In vivo 
survival and organ uptake of loaded carrier rat erythrocytes. / Biochem 120:286-291. 
40. Muzykantov VR, Zaltsman AB, Smirnov MD, Samokhin GP and Morgan BP (1996) 
Target-sensitive immunoerythrocytes: Interaction of biotinylated red blood cells 
with immobilized avidin induces their lysis by complement. Biochim Biophys Acta 
1279:137-143. 
41. Jordan JA, Alvarez FJ, Lotero LA, Herraez A, Diez JC and Tejedor MC (2001) In vitro 
phagocytosis of carrier mouse red blood cells is increased by Band 3 cross-linking or 
diamide treatment. Biotechnol Appl Biochem 34:143-149. 
42. Lotero LA, Jordan JA, Olmos G, Alvarez FJ, Tejedor MC and Diez JC (2001) Differential 
in vitro and in vivo behaviour of mouse ascorbate/Fe3+ and diamide oxidized 
erythrocytes. Biosci 21:857-871. 
43. Mishra PR and Jain NK (2000) Biotinylated methotrexate loaded erythrocytes for 
enhanced liver uptake. "A study on the rat". / Drug Targ 217. 
44. Kruse CA, Freehauf CL, Patel KR and Baldeschwieler JD (1987) Mouse erythrocyte 
carriers osmotically loaded with methotrexate. Biotechnol Appl Biochem 9:123-140. 
45. Mishra PR and Jain NK (2002) Biotinylated methotrexate loaded erythrocytes for 
enhanced liver uptake. "A study on the rat". Int} Pharm 145. 
46. Lotero LA, Olmos G and Diez JC (2003) Delivery to macrophages and toxic action of 
etoposide carried in mouse red blood cells. Biochim Biophys Acta 1620:160-166. 
Cells and Cell Ghosts as Drug Carriers 345 
47. Zocchi E, Tonetti M, Polvani C, Guida L, Benatti U and De Flora A (1988) In vivo liver 
and lung targeting of adriamycin encapsulated in glutaraldehyde-treated murine erythrocytes. 
Biotechnol Appl Biochem 10:555-562. 
48. Eichler HG, Gasic S, Bauer K, Korn A and Bacher S (1986a) In vivo clearance 
of antibody-sensitized human drug carrier erythrocytes. Clin Pharmacol Ther 40: 
300-303. 
49. Eichler HG, Rameis H, Bauer K, Korn A, Bacher S and Gasic S (1986b) Survival 
of gentamicin-loaded carrier erythrocytes in healthy human volunteers. Eur J Clin 
Invest! 16:39-42. 
50. Ogiso T, Iwaki M and Ohtori A (1985) Encapsulation of dexamethasone in rabbit 
erythrocytes, the disposition in circulation and anti-inflammatory effect. / Pharmacobiodyn 
8:1032-1040. 
51. Rossi L, Serafini S, Cenerini L, Picardi F, Bigi L, Panzani I and Magnani M (2001) 
Encapsulation of dexamethasone in rabbit erythrocytes, the disposition in circulation 
and anti-inflammatory effect. Biotechnol Appl Biochem 85. 
52. Noel-Hocquet S, Jabbouri S, Lazar S, Maunier JC, Guillaumet G and Ropars C (1992) 
Magnani M and DeLoach JR (eds.), The Use of Resealed Erythrocytes as Carriers and 
BioreactorsAdvances in Experimental Medicine and Biology 326, Plenum Press, New York, 
pp. 215-221. 
53. Kravtzoff R, Desbois I, Lamagnere JP, Muh JP, Valat C, Chassaigne M, Colomba P and 
Ropars M (1996) Improved pharmacodynamics of L-asparaginase-loaded in human 
red blood cells. Eur J Clin Pharmacol 49:465^70. 
54. Rossi L, Bianchi M and Magnani M (1992) Increased glucose metabolism by enzymeloaded 
erythrocytes in vitro and in vivo normalization of hyperglycemia in diabetic 
mice. Biotechnol Appl Biochem 15:207-216. 
55. Lizano C, Sanz S, Luque J and Pinilla M (1998) In vitro study of alcohol dehydrogenase 
and acetaldehyde dehydrogenase encapsulated into human erythrocytes by an 
electroporation procedure. Biochim Biophys Acta 1425:328-336. 
56. Magnani M, Laguerre M, Rossi L, Bianchi M, Ninfali P, Mangani F and Ropars C (1989) 
Acetaldehyde dehydrogenase-loaded erythrocytes as bioreactors for the removal of 
blood acetaldehyde. Alcohol Clin Exp Res 13:849-859. 
57. Magnani M, Fazi A, Magnani F, Rossi L and Mancini U (1993) Methanol detoxification 
by enzyme-loaded erythrocytes. Biotechnol Biochem Appl 18:217-226. 
58. Sanz S, Lizano C, Luque J and Pinilla M (1999) In vitro and in vivo study of glutamate 
dehydrogenase encapsulated into mouse erythrocytes by hypotonic dialysis procedure. 
Life Sci 65:2781-2789. 
59. Magnani M, Mancini U, Bianchi M and Fazi A (1992) Magnani M and DeLoach JR 
(eds.), The Use of Resealed Erythrocytes as Carriers and Bioreactors. Advances in Experimental 
Medicine and Biology 326, Plenum Press: New York, pp. 189. 
60. Ito Y, Ogiso T, Iwaki M and Atago H (1987) Encapsulation of human urokinase in rabbit 
erythrocytes and its disposition in the circulation system in rabbits. / Pharmacobiodyn 
10:550-556. 
346 Lanao & Sayalero 
61. Garin M, Rossi L, Luque J and Magnani M (1995) Lactate catabolism by enzyme-loaded 
red blood cells. Biotechnol Appl Biochem 22:295-303. 
62. Adriaenssens K, Karcher D, Marescau B, Van Broeckhoven A and Terheggen HC (1984) 
Hyperargininemia: The rat as a model for the human disease and the comparative 
response to enzyme replacement therapy with free arginase and arginase loaded erythrocytes 
in vivo. Int J Biochem 16:779-786. 
63. Petrikovics I, Pei L, McGuinn WD, Cannon EP and Way JL (1994) Encapsulation 
of rhodanese and organic thiosulfonates by mouse erythrocytes. Fundam Appl 
Toxicol 23:70-75. 
64. Pei L, Omburo G, McGuinn WD, Petrikovics I, Dave K, Raushel FM, Wild JR, DeLoach 
JR and Way JL (1994) Encapsulation of phosphotriesterase within murine erythrocytes. 
Toxicol Appl Pharmacol 124:296-301. 
65. Bustos NL and Batlle AM (1989) Enzyme replacement therapy in porphyrias: V. 
In vivo correction of delta-aminolaevulinate dehydratase defective in erythrocytes in 
lead intoxicated animals by enzyme-loaded red blood cell ghosts. Drug Des Del 5: 
125-131. 
66. Hamarat Baysal S and Uslan AH (2002) In vitro study of urease/ AlaDH enzyme system 
encapsulated into human erythrocytes and research into its medical applications. Artif 
Cells Blood Substit Immobil Biotechnol 30:71-77. 
67. Bax BE, Bain MD, Fairbanks LD, Simmonds HA, Webster AD and Chalmers RA (2000) 
Carrier erythrocyte entrapped adenosine deaminase therapy in adenosine deaminase 
deficiency. Adv Exp Med Biol 486:47-50. 
68. Flynn G, Hackett TJ, McHale L and McHale AP (1994) Encapsulation of the thrombolytic 
enzyme, brinase, in photosensitized erythrocytes: A novel thrombolytic system based 
on photodynamic activation. J Photochem Photobiol B Biol 76:193-196. 
69. Bax BE, Bain MD, Ward CP, Fensom AH and Chalmers RA (1996) The entrapment of 
mannose-terminated glucocerebrosidase (alglucerase) in human carrier erythrocytes. 
Biochem Soc Trans 24:442S. 
70. Oettgen HF, Old LJ, Boyse EA, Campbell HA, Philips FS, Clarkson BD, Tallal L, 
Leeper RD, Schwartz MK and Kim JH (1967) Inhibition of leukaemia in man by 
L-asparaginase. Cancer Res 27:2619-2631. 
71. Fraternale A, Casabianca A, Tonelli A, Chiarantini L, Brandi G and Magnani M (2001) 
New drug combinations for the treatment of murine AIDS and macrophage protection. 
Eur J Clin Investig 31:248-252. 
72. Magnani M, Rossi L, Fraternale A, Silvotti L, Quintavalla F, Piedimonte G, Matteucci D, 
Baldinotti F and Bendinelli M (1995) FIV infection of macrophages: In vitro and in vivo 
inhibition by dideoxycytidine 5'-triphosphate. Vet Immunol Immunopathol 46:151-158. 
73. Magnani M, Bianchi M, Rossi L and Stocchi V (1989) Human red blood cells as bioreactors 
for the release of 2',3'-dideoxycytidine, an inhibitor of HIV infectivity. Biochem 
Biophys Res Commun 164:446-457. 
74. Fraternale A, Casabianca A, Rossi L, Chiarantini L, Schiavano GF, Palamara AT, Garaci E 
and Magnani M (2003) Erythrocytes as carriers of reduced glutathione (GSH) in the 
treatment of retroviral infections. / Antimicrob Chemother 52:551-554. 
Cells and Cell Ghosts as Drug Carriers 347 
75. Fraternale A, Rossi L and Magnani M (1996) Encapsulation, metabolism and release of 
2-fluoro-ara-AMP from human erythrocytes. Biochim Biophys Acta 1291:149-154. 
76. Al-Achi A, Greenwood R and Walker B (1994) Buccal administration of erythrocyteghosts-
insulin in rats. / Control Rel 267. 
77. Nielsen PE (2000) Antisense peptide nucleic acids. Curr Opin Mol Ther 2:282-287. 
78. Chiarantini L, Cerasi A, Fraternale A, Andreoni F, Scari S, Giovine M, Clavarino E and 
Magnani M (2002) Inhibition of macrophage iNOS by selective targeting of antisense 
PNA. Biochemistry 41:8471-8477. 
79. Garin MI, Lopez RM and Luque J (1997) Pharmacokinetic properties and in vivo biological 
activity of recombinant human erythropoietin encapsulated in red blood cells. 
Cytokine 9:66-71. 
80. Eichler HG, Schneider W, Raberger G, Bacher S and Pabinger I (1986) Erythrocytes as 
carriers for heparin. Preliminary in vitro and animal studies. Res Exp Med 186:407-412. 
81. Feder R, Nehushtai R and Mor A (2001) Affinity driven molecular transfer from erythrocyte 
membrane to target cells. Peptides 22:1683-1690. 
82. Olmos G, Lotero LA, Tejedor MC and Diez JC (2000) Delivery to macrophages of interleukin 
3 loaded in mouse erythrocytes. Biosci Rep 70:399-410. 
83. Polvani C, Gasparini A, Benatti U, De Flora A, Silvestri S, Volpini G and Nencioni L 
(1991) Murine red blood cells as efficient carriers of three bacterial antigens for the 
production of specific and neutralizing antibodies. Biotechnol Appl Biochem 14:347-356. 
84. Garin MI, Lopez RM, Sanz S, Pinilla M and Luque J (1996) Erythrocytes as carriers for 
recombinant human erythropoietin. Pharm Res 13:869-874. 
85. Feder R, Nehushtai R and Mor A (2001) Affinity driven molecular transfer from erythrocyte 
membrane to target cells. Peptides 22:1683-1690. 
86. Lejeune A, Moorjani M, Gicquaud C, Lacroix J, Poyet P and Gaudreault R (1994) 
Nanoerythrosome, a new derivative of erythrocyte ghost: Preparation and antineoplastic 
potential as drug carrier for daunorubicin. Anticancer Res 14:915-919. 
87. Moorjani M, Lejeune A, Gicquaud C, Lacroix J, Poyet P and Gaudreault RC (1996) 
Nanoerythrosomes, a new derivative of erythrocyte ghost II: Identification of the mechanism 
of action. Anticancer Res 2831. 
88. Lejeune A, Poyet P, Gaudreault RC and Gicquaud C (1997) Nanoerythrosomes, a new 
derivative of erythrocyte ghost: III. Is phagocytosis involved in the mechanism of 
action? Anticancer Res 3599. 
89. Desilets J, Lejeune A, Mercer J and Gicquaud C (2001) Nanoerythrosomes, a new derivative 
of erythrocyte ghost: IV. Fate of reinjected nanoerythrosomes. Anticancer Res 1741. 
90. Mishra PR and Jain NK (2000) Reverse biomembrane vesicles for effective controlled 
delivery of doxorubicin HC1. Drug Del 7:155-159. 
91. El-Aneed A (2004) An overview of current delivery systems in cancer gene therapy. 
/ Control Rel 94:1-14. 
92. Klink D, Schindelhauer D, Laner A, Tucker T, Bebok Z, Schwiebert EM, Boyd AC and 
Scholte BJ (2004) Gene delivery systems-gene therapy vectors for cystic fibrosis. / Cyst 
Fibros 203. 
348 Lanao & Sayalero 
93. Tomanin R and Scarpa M (2004) Why do we need new gene therapy viral vectors? 
Characteristics, limitations and future perspectives of viral vector transduction. Curr 
Gene Ther 4:357-372. 
94. Studeny M, Marini FC, Dembinski JL, Zompetta C, Cabreira-Hansen M, Bekele BN, 
Champlin RE and Andreeff M (2004) Mesenchymal stem cells: Potential precursors for 
tumor stroma and targeted-delivery vehicles for anticancer agents. / Natl Cancer Inst 
96:1593-1603. 
95. Ehtesham M, Kabos P, Kabosova A, Neuman T, Black KL and Yu JS (2002) The use 
of interleukin 12-secreting neural stem cells for the treatment of intracranial glioma. 
Cancer Res 62:5657-5663. 
96. Su L, Lee R, Bonyhadi M, Matsuzaki H, Forestell S, Escaich S, Bohnlein E and 
Kaneshima H (1997) Hematopoietic stem cell-based gene therapy for acquired immunodeficiency 
syndrome: Efficient transduction and expression of RevMlO in myeloid 
cells in vivo and in vitro. Blood 89:2283-2290. 
97. Sunkomat JN and Gaballa MA (2003) Stem cell therapy in ischemic heart disease. Cardiovasc 
Drug Rev 21:327-342. 
98. Huber A, Padrun V, Deglon N, Aebischer P, Mohler H and Boison D (2001) Grafts of 
adenosine-releasing cells suppress seizures in kindling epilepsy. Proc Natl Acad Sci USA 
98:7611-7616. 
99. Boison D, Huber A, Padrun V, Deglon N, Aebischer P and Mohler H (2002) Seizure 
suppression by adenosine-releasing cells is independent of seizure frequency. Epilepsia 
43:788-796. 
100. Shao J, DeHaven J, Lamm D, Weissman DN, Malanga CJ, Rojanasakul Y and Ma JK 
(2001b) A cell-based drug delivery system for lung targeting: II. Therapeutic activities 
on B16-F10 melanoma in mouse lungs. Drug Del 8:71-76. 
101. Alby L and Auerbach R (1984) Differential adhesion of tumor cells to capillary endothelial 
cells in vitro. Proc Natl Acad Sci USA 81:5739-5743. 
16 
Cochleates as Nanoparticular Drug 
Carriers 
Leila Zarif 
1. Introduction 
In spite of the availability of many non-traditional novel dosage forms, oral route 
remains the most attractive way for administration of therapeutical materials. 
However, many therapeutic agents, especially the increasing number of biological 
molecules cannot be taken up by intestine due to their intrinsic impermeability 
to tissue membranes and the enzymatic degradation through the wall of the GI tract. 
Carrier systems that facilitate intestine uptake of these molecules are of major interests 
in the drug delivery arena. Moreover, drug delivery systems that provide a route 
of administration that does not involve injection can improve patient compliance 
and expand the market for existing, injectable, drugs. The factors which are important 
for the oral efficiency of a vehicle system have been repeatedly summarized in 
the literature.1,2 Small particle size, appropriate surface properties, mucoadhesive 
and targeting moieties, stability, as well as dose are the major factors imparting the 
efficiency of oral uptake. 
Producing formulations of poorly soluble drugs with high bioavailability is 
an even higher challenge. Known technologies are nanocrystals and nanoparticles 
which use the approach of enhancing the bioavailability by a decrease in particle 
size, resulting in an increase of surface area and subsequently a faster dissolution. 
Other technologies such as solid dispersions, polymeric micelles and selfemulsifying 
systems were developed to increase the drug solubility. 
349 
350 Zarif 
Many lipid-based systems were developed to enhance oral bioavailability3,4 
Examples are lipid-based emulsions & microemulsions5-7; Solid lipid nanoparticles 
(SLN), a high melting point lipids enclosed in a surfactant layer8,9 adequate 
to enhance the oral bioavailability of poorly absorbed drugs; Lipid nanocapsules 
(LNC) for oral, injectable use10 and improved bioavailability11; Lipid nanospheres 
prepared from egg lecithin and soybean, described for their low toxicity12 and 
higher efficacy, compared with other delivery systems when incorporating amphotericin 
B,13 due to their smaller particle size and lower uptake by reticuloendothelial 
system.14,15 Recently, solid lipid microparticles, prepared by the solvent-in-wateremulsion-
diffusion technique, were described for the encapsulation and oral delivery 
of insulin.16 
In particular, lipid-based cochleate delivery system appears to provide answers 
to oral delivery challenges by (1) formulating different kind of molecules, especially 
hydrophobic ones17,18 and (2) protecting the sensitive and biologically active 
molecules from harsh environmental conditions. 
In this review, we will focus on cochleates nanoparticular drug carrier and will 
present the main features and the state of the art of this delivery technology. 
2. Cochleates Nanoparticles in Oral Delivery 
2.1. Cochleate structure 
Cochleates were first described by Dimitrious Papahadjopoulos and his co-workers 
in 1975 as precipitates formed by the interaction of negatively charged phosphatidylserine 
and calcium.19-21 He named these cylindrical structures "cochleate", 
meaning shell in the Greek language because of their rolled-up form, and explained 
the mechanism of cochleates formation by the fusion of negatively charged vesicles 
induced by the calcium cation22 (Fig. 1). 
These cigar-like structures have gained interest as antigen delivery system for 
vaccine applications.23 More recently cochleates were studied as tools to deliver 
small molecule drugs.17,18,24 A cochleate lipid formulation of amphotericin B has 
been developed as an oral composition to treat systemic fungal infections.24-26 
Other medical and non-medical applications are also under investigation.27 
2.2. Cochleate preparation 
2.2 A. Which phospholipid and which cation to use? 
Cochleates are a phospholipid-ion precipitates. Does that mean that cochleate is a 
structure obtained from precipitation of any phospholipid with any ion as presented 
in some litterature?,28 i.e. a complex of negatively charged phospholipid with any 
cation or a complex made from a positively charged lipid with any anion? 
Cochleates as Nanoparticular Drug Carriers 3 51 
f) Ca 0 fusion ( \ I 1 ///EPTA. 
A B C D E F 
Fig. 1. Cochleate cylindrical structure and mechanism of formation (adapted from Refs. 19 
and 69 with permission). 
Papahadjopoulos has given in 1975 this appellation to a rolled phospholipid 
structure. So far, to our knowledge no physico-chemical evidence on the obtention 
of such cigar-like structure from positively charged phospholipid with an anion 
had been described; on the contrary, extensive litterature is available on obtaining 
these cigar-like structure when negatively charged phospholipid such as phosphosphatidylserine 
(PS) had been precipitated with a cation such as calcium.17'18'20-22'29-30 
Other negatively charged phospholipids, such as phosphatidic acid (PA) or phosphatidyl 
glycerol derivatives, have been studied as well. Mixture of negatively 
charged phospholipids with other lipids can lead to cochleate formation. In this 
case, the cochleate formation depends on the negatively charged lipid/other lipid 
ratio and depends on the nature of the negatively charged lipid in the mixed lipid 
system. For example, PA derivatives form cochleate domains after the addition of 
calcium cation. However, when mixed with the corresponding diacylphosphatidylcholine 
(PC) and diacylphosphatidylethanolamine (PE), it was found that up to 
20 mole% of PC or PE can be introduced into the cochleate phase of PA(Ca2+), 
above which a distinct PC rich or PE-rich phase appears.31 
Other phospholipid derivatives such as galactosphingolipid hydroxy fatty acid 
cerebroside were reported to form cochleate cylinders by thermal mechanical treatment 
of glycol suspensions.32 However, the addition of conjugated lipid, such as 
352 Zarif 
poly(ethylene glycol)-lipid conjugates to PS vesicles, inhibited the calcium-induced 
fusion.33 
In general, an additional desired feature of an oral drug delivery system is 
that the excipient permitting this transport to be classified is generally regarded 
as safe (GRAS). Soy phosphatidylserine fits this criteria. Furthermore, Soy PS 
has been used as a nutrient supplement since early 1980s. Clinical trials showed 
that PS may play a role in supporting mental functions in aging brains such as 
enhancing the memory, improving learning ability,34-41 reducing the stress42'43 and 
anxiety.44 
Cochleates can be made from purified soy phosphatidylserine, which represents 
an affordable source of raw material.45 A study comparing the purified soy 
phosphatidylserine (PSPS) to non-purified soy PS (NPSPS) has been disclosed in 
this patent, showing that PS should be present in an amount of at least 75% of the 
total lipid in order to allow the formation of cochleates. The other 25% phospholipids 
present can be selected either from the anionic group such as phosphatidic 
acid, phosphatidylglycerol, phosphatidyl inositol or phosphatidylcholine. PSPS 
cochleates can be loaded with different bioactive materials such as nutritional supplement, 
vitamins, antiviral, antifungal, small peptides. Proof of principle of the 
use of purified soy PS has been achieved using a polyene antifungal agent, amphotericin 
B. The preparation method for amphotericin B cochleates can be either via 
High pH-trapping or film method18 or by hydrogel method;29 the latter leading to 
nanocochleates formation. 
The nature of the cation is an important factor in cochleate formation. In the 
precipitation process, divalent cations are preferred to monovalent cations. Monovalent 
cations such as Na+ were described to prevent the cochleate formation.46 
Increases concentration of Na+ ions was shown to interfere with the destabilization 
effect of Ca2+. A critical Ca/PS ratio is necessary for the destabilization effect 
of divalent cations and the formation of cochleate phases.46 
The formation of cochleate is easier from small unilamellar vesicles (SUV). 
However, multilamellar vesicles (MLV) can also lead to cochleate formation. In this 
case, the first mechanism is a destabilization of the outer bilayer of PS by Ca2+ 
which causes its collapse, leading to a higher access of Ca2+ to inner PS bilayers 
and so forth. 
2.2.2. Which molecules can be entrapped in cochleates nanoparticles 
Due to the intrinsic nature of the lipid-contained cochleates, these nanoparticles 
can encapsulate a variety of molecules of all shapes and sizes. Preference is given, 
however, to hydrophobic molecules, for which a need to enhance chemical stability 
or bioavailability is desired [Fig. 2(a)]. Amphiphatic molecules which can easily 
Cochleates as Nanoparticular Drug Carriers 353 
Fig. 2. Type of molecules which can be encapsulated into lipid based cochleate (adapted 
from Ref. 18 with permission). 
insert in the membrane bilayers [Fig. 2(b)], negatively charged moiety [Fig. 2(c)] 
or positively charged moiety [Fig. 2(d)] could be encapsulated in the cochleate 
nanoparticle structure. 
The nature of the drug influence the percentage of encapsulation. Hydrophobic 
drag shows a quantitative encapsulation, whereas less was seen for amphiphatic 
molecules. For instance, doxorubicin which presents hydrophobic regions is a 
water-soluble drug, has a partition between the bilayers and the external aqueous 
phase [Fig. 2(b)]. As calcium induces dehydration of the interbilayer domains, 
the amount of water in this region is low,47 therefore, small hydrophilic molecules 
will not be suitable for cochleate system. 
2.2.3. Multiple ways of preparing cochleates 
Several processes were developed to obtain cochleates with a nanosize range, with 
the objective to allow oral delivery.24,29'48-59 Particle size is process dependent. When 
a small nanosized particle is desired, the "hydrogel method" can be used, based 
on the use of an aqueous-aqueous emulsion system.29 Briefly, this method consists 
of 2 steps: The preparation of small size liposomes either by high pH method18'25 
or by film method,18 then the liposomes are mixed with a high viscosity polymer 
354 Zarif 
such as dextran. The dextran/liposome phase is then injected into a second, nonmiscible, 
polymer (i.e. PEG). The calcium was then added and diffused slowly from 
one phase to another, resulting in the formation of nanocochleates. The final step 
is the washing of the gel. These nanosized cochleates showed potential in the oral 
delivery of drugs.18,29,48,59 
Electron microscopy and X-Ray crystallography of the nanoparticles show a 
unique multilayered structure consisting of continuous, solid lipid bilayer sheets, 
rolled up in a spiral with no internal aqueous space and the localization of AmB in 
the lipid bilayer.25 
Other preparation techniques are known, e.g. the trapping method, useful for 
the encapsulation of hydrophilic and hydrophobic molecules,17'18 which consist 
in the preparation of the liposomal suspension containing the drug either in the 
aqueous space of liposome (when hydrophilic) or intercalated in between the bilayers 
(when hydrophobic). A step of addition of calcium follows, and an aggregate 
of cochleates are formed. The cochleates made by the Trapping method present 
higher aggregation compared with other methods. This has been demonstrated 
using Electron microscopy after Freeze-fracture.25 
Another method was developed for hydrophobic drugs,61 known as "the solvent 
drip method" which consists of preparing a liposomal suspension separately 
based on soy PS and a hydrophobic or amphipathic cargo moiety solution. Solvent 
for hydrophobic drug can be selected from DMSO, DMF. The solution is then added 
to liposomal suspension. Since the solvent is miscible in water, a decrease of the 
solubility of the cargo moiety is observed, which associates at least in part with the 
lipid-hydrophobic liposomal bilayers. The cochleates are then obtained by addition 
of calcium and the excess solvent is being washed. 
Usually, the cochleate formation can be characterized by optical microscopy 
when they are present in needle form in the micrometer size range. In this case, 
direct observation using a higher magnification can be used.25 When nanocochleate 
are obtained, optical microscope can be used as an indirect method to assess the 
formation of cochleate, i.e. observation of the liposome formation after chelation 
of the calcium present, by addition of EDTA (ethylene diamine tetraacetate) to 
nanocochleate. A more sophisticated method is the electron microscopy after freezefracture18'
25 which allows the observation of the tighted packed bilayers. Recently, 
other methods were described using Laurdan (6-dodecanoyl-2-dimethylamino 
naphtalene) to monitor the cochleate phase formation.62 In this case, the lipid vesicles 
are labeled with Laurdan and the addition of calcium to the laurdan labeled 
vesicles resulted in a shift in the emission peak maximum of Laurdan. Due to 
dipolar relaxation, excitation and emission, generalized polarization (GPgx and 
GPEm) indicates the transition from a LC to a rigid and dehydrated cochleate 
phase. 
Cochleates as Nanoparticular Drug Carriers 355 
2.3. Cochleates as oral delivery system for antifungal agent, 
amphotericin B 
Among the drug of choice using nanocochleate delivery system, amphotericin B 
(AmB) presented all aspects of a good candidate. Amphotericin B is a hydrophobic 
drug with poor oral bioavailability. This drug had been used for decades in 
injectable form to treat systemic fungal infections of Candida, cryptococcus and 
aspergillosis species.63-65 
Lipid formulations of Amphotericin B such as liposomes, lipid complexes, lipid 
emulsions and colloidal dispersions, were developed with the aim to achieve a 
higher therapeutic index.26-66 These formulations indeed showed enhanced therapeutic 
index, even though none of these formulations showed ability to deliver 
AmB orally. Cocheate technology seems to offer the advantage over other delivery 
systems in providing the possibility for the oral delivery of AmB. Oral administration 
of amphotericin B cochleates (CAMB) to healthy mice achieved potentially 
therapeutic concentrations in key target tissues.51 
Preclinical studies demonstrate a promising activity of CAMB in murine 
models of clinically relevant invasive fungal infections such as disseminated 
candidiasis,25'48,67 disseminated aspergillosis17,18-58'59 and central nervous system 
cryptococcosis.68 
2.3.1. In candidiasis animal model 
In Candida albicans infected murine animal model, AmB cochleates showed potential 
either after intraperitoneal (i.p.) or oral (p.o.) administration.17,18,48,49,54,55,57,60,66-68 
After i.p. administration CAMB provided protection against C. albicans at doses 
as low as 0.1 mg/kg/day, kidney tissues burden showed that CAMB was more 
potent than Fungizone® at 1 mg/kg/day and was equivalent to AmBisome® at 
10 mg/kg/day18,25,60 (Fig. 3). CAMB was also effective after oral administration. 
Complete eradication of C. albicans from the lungs was noticed after p.o. administration 
at 2.5 mg/kg/day. These results were comparable to i.p. Fungizone® at 
2.0 mg/kg/day.48,54-56 
2.3.2. In aspergillosis animal model 
Oral administration of CAMB was shown to be protective in a dosedependent 
manner against systemic infection of Aspergillus fumigatus in animals 
immunosusppressed with cyclophosphamide.58,59 In this mouse model, intragastric 
administration of CAMB at 40 mg/kg/day for 15 days resulted in 80% survival, 
while Fungizone at 4 mg/kg/day (i.p.) resulted in 20% survival; higher doses of 
Fungizone were lethal to animals. 
356 Zarif 
0) 
•J 
t/1 
UJ 
0) 
~-~ :-> Uo 
1071 
106 - 
105 - 
104 - 
mJ- 
102- 
1 0 ' • 
10°- 
1 U 1 1 . 1 1 1 1 1 • > 
Control 0.1 1.0 10.0 0.1 1.0 10.0 0.1 1.0 
AmB Dose Concentration (mg/kg) 
Fig. 3. Kidneys tissue burden of infected mice treated with either CAMB (•), Fungizone 
(•) or AmBisome (•), compared with controls (T) (from Ref. 18 with permission) 
2 
trt 
• 1 1 I J ffl -1 u 
* 
1 i* 
* JL 
' l 5 ! 
U JJ M • M. , , 
• Liver 
• Kidney 
• 1.urn's 
1 
control DAMB Smg/kg lOmg/kg 20mg/kg 30mg/kg 40mg/kg 
5- 
a a c; 
& 
<=> 
K 
3 
<5 
-J 
(3 
CIO 
s? ~ — • < 
Sfl 
40 
30 
20 
10 
Concentration of Drug 
Fig. 4. Tissue burden for mice infected in a model of invasive aspergillosis after oral administration 
of CAMB (from Ref. 58 with permission). 
The tissue fungal burden for target organs, kidneys, liver and lungs, demonstrated 
the benefic effect of CAMB (Fig. 4 ). CAMB showed a pronounced dosedependent 
reduction in the fungal burden in all organs. The near eradication 
of Aspergillus was observed above a concentration of 20mg/kg/day. CAMB at 
30 mg/kg (PO) was as effective as CAMB at 20 mg/kg (PO) in reducing fungal 
tissue burden.58 
n? 
o<> 
00° 
... . t 
I 
Cochleates as Nanoparticular Drug Carriers 357 
2.3.3. In cryptococcal meningitis animal model 
Oral amphotericin B cochleates were effective in a murine cryptococcal meningitis 
model with an 80% survival after 17 days, obtained after oral treatment 
with CAMB (lOmg/kg) to mice having intracerebral infection with cryptococcus 
neoformans.68 
2.3.4. Toxicity of amphotericin B cochleates 
In vitro, Amphotericin B cochleates (CAMB) showed a low toxicity on red blood cells 
when compared with Fungizone (DAMB). CAMB showed no hemoglobin release 
and therefore no hemolysis of red blood cells when incubated at 500 ^g/ml. In 
contrast, DAMB was hemolytic at 10 /xg/ml due to the presence of the detergent, 
sodium desoxycholate.25 
In vivo, CAMB was non toxic to mice when administered orally at 
50mg/kg/day for 14 days. No nephrotoxicity was observed as demonstrated by 
the normal BUN level, and the histopathology of kidneys, lungs, liver, spleen and 
GI tract showed that animals dosed with CAMB were comparable to controls.18 
2.3.5. Pharmacokinetics of amphotericin B cochleates 
Oral pharmacokinetics?)^ 
Pharmacokinetic studies have shown that after oral administration of CAMB, AmB 
is distributed into the target tissues (e.g. brain, liver, lung, spleen and kidneys)18,50'52 
in healthy mice and AmB tissue level suggests a zero-order uptake process for all 
tissues. 
When CAMB was administered po to C57BL/6 mice at lOmg/kg (n = 5), 
and blood and tissues collected and AmB level measured by HPLC, blood 
shows a plateau-shaped profile with Tmax = 6h and Cmax = 0.05mg/ml. Noncompartmental 
(NCA) analysis showed blood AUC0-oo = 1.20/xg*h/ml, ti/2 = 
12.8 h, MRTo_oo = 21.1 h, Cl/F = 139.2ml/min/kg, Vz /F = 153.91 L/kg. AmB tissue 
exposure (AUCo-oo, .ig*h/g) evaluated using NCA was greater for lungs (23.11), 
followed by liver (16.91), spleen (15.40) kidneys (14.97) and heart (3.34). Tissue elution 
ti/2(h): kidneys 9.3, lungs 5.6, heart 5.3, liver 4.9 and spleen 4.3. For all tissues, 
Tmax = 12 h and Cmax ranged between 0.23/zg/ml for heart and 1.58/xg/ml for 
lungs.52 
The delivery of AmB by cochleates after multiple oral doses (10) was assessed 
in the same mouse model and was compared with AmBisome. It was found 
that cochleate provides therapeutic levels in tissue and presents better delivery 
and transfer efficiency of AmB to the target tissue, as well as better tissue 
penetration.53 
358 Zarif 
The ability of cochleate vehicles to deliver systemic AmB after single or multiple 
oral dosing suggest the potential of CAMB formulations to treat and prevent 
systemic fungal infections. 
Pharmacokinetics 
AmB given intraveneously (IV) to mice showed a two-phase pharmacokinetic 
profile.69,70 Pharmacokinetic analysis in target tissues (liver, spleen, kidney and 
lungs) shows a multi-peak profile, large AUC and MRT. 
After IV administration of 0.625 mg/kg, AMB presented a two-phase blood 
concentration time course [Fig. 5(A)]. This profile is characterized by a very fast 
distribution phase and an elimination phase with t1/2 = 11.68 hrs. The AUCo-oo w a s 
1.006 A<,g*h/ml, CI = 10.36 ml/min/kg, MRT0_oo = 15.41 hrs and Vs s = 9.587 L/kg. 
This pharmacokinetic profile indicates that CAMB is removed fast from blood. 
In addition, the large Vss also indicates a large distribution into the tissues. The 
results obtained in target tissues showed this extensive distribution and penetration 
[Fig. 5(B)]. 
Calculation of pharmacokinetic parameters showed that the main target tissues 
have a large AMB exposure reflected in the AUC and CMAX values (Table 1), as well 
as the tissue to blood AUC ratio. 
The large AMB exposure in liver and spleen suggests involvement of the 
mononuclear phagocyte system (MPS) in the removal of CAMB. Cochleates are 
particulates that can be quickly cleared from the circulation by the macrophages of 
the reticular endothelial system (RES) related to the liver and the spleen. In addition, 
"physical retention" seems to play a role in the kinetic profile of the lungs due 
to its capillary nature. 
Time (hrs) "* 
0 10 20 30 40 50 
Time (hours) 
Fig. 5. (A) AMB profile in blood after a single dose (B) IV PK profile of AMB in target 
tissues, (from Ref. 69, with permission). 
Cochleates as Nanoparticular Drug Carriers 359 
Table 1 Pharmacokinetics parameters for CAMB in different 
target organs after IV administration to C57BL/6 
mice (n = 5 per time point) (From Ref. 69, with 
permission). 
Tissue3 
Liver 
Spleen 
Lung 
Kidney 
Heart 
Intestine 
Stomach 
AUCo-oo 
(/ig*h/g) 
474.519 
116.388 
39.707 
12.564 
0.970 
9.173 
8.184 
T max 
(min) 
10 
2 
2 
5 
5 
20 
20 
*~ max 
(Mg/g) 
8.559 
6.633 
16.408 
1.032 
0.478 
0.609 
0.343 
t l/2>-z
b 
(hrs) 
75.03 
66.71 
22.34 
21.86 
2.82 
13.88 
20.77 
This phenomenon and the mobility of the macrophages seem to cause certain 
redistribution of cochleates that gives a multi-peak and plateau shape profiles in 
liver and spleen. Finally, AMB was also detected in bile and intestine contents, 
suggesting that bile excretion may be an additional elimination route. 
2.4. Other potential applications for cochleates 
2.4.1. Cochieate for the delivery of antibiotics 
As cochieate has shown a high affinity to be engulfed by macrophages [Fig. 6(A)] 
probably due to a dual mechanism, the cochieate essential particulate feature71 and 
possibly a PS receptor mediated internalization of the cochieate into macrophage.72 
Fig. 6. Uptake of amphotericin B cochleates by J774 macrophages as seen by (A) fluorescence 
microscopy, (B) confocal microscopy (from Ref. 17, with permission). 
360 Zarif 
This particulate system would have potential for the delivery of antibacterial 
agents such as aminoglycosides and vancomycin.17 Illustration is given by the 
encapsulation of clofazimine, an anti-TB drug, and tobramycin, an aminoglycoside 
antibiotic used in treating bacterial infections, both given intraveneously thus far. 
The cochleate system may possibly offer a new oral way of delivery. 
2.4.2. Delivery of clofazimine 
Clofazimine cochleates were prepared by the Trapping method.18 Clofazimine 
is a known hydrophobic anti-TB drug, the efficacy of Clofazimine cochleate 
was assessed by measuring the IC50 in Vero Cells and in bone marrow derived 
macrophage (BM-M).73 Clofazimine cochleates exhibit a greater decrease in toxicity 
versus free clofazimine and had a higher efficacy in killing intracellular 
M. Tuberculosis than free clofazimine:2 Log reduction (CE99) was achieved at 
20.9 /xg/ml for cochleates, while free clofazimine was toxic at this concentration. 
This shows that encapsulation of clofazimine in cochleates potentiates the antimicrobial 
efficacy of the drug, i.e. when higher concentration of drug can be used 
because of less toxicity, bactericidal levels of the drug could be attained. 
2.4.3. Delivery of tobramycin 
A recent research work has been published on the possible use of nanocochleates as 
an oral delivery system for Tobramycin.74 Tobramycin is a well known aminoglycoside 
antibiotic used in treating bacterial infections, and is usually administered by 
intravenous (i.v.) infusion, intramuscular (i.m.) injection, or inhalation. This aminogycoside 
drug is known for its side effects such as mineral depletion (i.e. calcium, 
magnesium, potassium) after i.v. administration.75,76 
In this work, the author described that tobramycin which is positively charged 
at low pH, will be encapsulated in the inter-bilayer space of cochleates. The fusion 
of unilamellar liposomes is no longer induced by a metal cation such as Ca2+, 
but by the organic molecule to be encapsulated. The cochleate cylinders formation 
has been described by Papahadjoupolos as resulting partly from the intrinsic 
properties of the calcium cation. Indeed, phosphatidylserine shows considerable 
selectivity for calcium due to the propensity of calcium to lose part of its hydration 
shell, and to displace water upon complex formation.19'77 In the cochleate solid 
crystalline structures formation, calcium plays a crucial role in bringing bilayers 
together closely through partial dehydration of the membrane surface and the crosslinking 
of opposing molecules of phosphatidylserine. In our opinion, in this recent 
work where formation of cochleate is claimed with no calcium present, additional 
Cochleates as Nanoparticular Drug Carriers 361 
relevant physico-chemical evidence on cochleate formation and the localization of 
the drug in the interbilayer space will be needed. 
2.4.4. Cochleate for the delivery of anti-inflammatory drugs 
As a result of the deep embedding of the molecules in the cochleates structures, 
drug molecules are hidden from the outside environment. This should have two 
beneficial effects: one is to hide and protect the molecule from the degradation due 
to environment; the other is to protect, the environment when needed, from the 
active molecule when such molecule presents side effects. 
This is the case of anti-inflammatory drugs, which associates cure to the disturbance 
of GI tract (stomach for instance). Cochleates were described to act beneficially 
in this area, reducing the stomach irritation when anti-inflammatory drugs 
such as aspirin is hidden in the cochleate structure, and administered to a carrageenan 
rat model for acute inflammation.27,61 
2.5. Othet uses of cochlea tes 
Cochleates were also described as vehicles for nutrients27 as an improved drug 
and contrast agent delivery system,28 as well as intermediate in the preparation of 
special liposomes such as Large Unilamellar Vesicles (LUV) and proteoliposomes. 
In fact, the discovery of the cochleate structures was a result of the desire to prepare 
LUV by Pr papahadjoupoulos,19'20 which were developed for the delivery of 
hydrophilic drugs. Proteoliposomes prepared from cochleates intermediates were 
described for vaccine applications in general,78 and more recently, when containing 
lipopolysaccharide as a novel adjuvant.79 
3. Conclusion 
Cochleates lipid-based nanocarrier appears to have potential for the oral delivery 
of bioactive molecules. Future work should be directed towards more fundamental 
science, as many research aspects of the cochleate drug carrier system are still hardly 
known (e.g. localization of the drug in lipid bilayers, impact of multivalent cations 
on the cochleate formation, mechanism of action of cochleate after oral uptake). In 
addition, the development of friendly analytical assays to monitor the drug localization 
and loading percentage in cochleates will be desired. This nano drug carrier 
is currently under development by Biodelivery Sciences International.27 Having 
the first drug-cochleate in the market place represents a big challenge. For instance, 
when oral amphotericin B cochleates are ultimately available for patients, thus will 
provide a new opening in the treatment of systemic fungal infections. 
362 Zarif 
References 
1. Chien YW (1992) Novel drug delivery systems. Drugs and the Pharmaceutical Sciences, 
Vol. 50. Marcel Dekker: New York, NY. 
2. Rathbone MJ, Hadgraft } and Michael SR (2003) Modified-release drug delivery technology. 
Drugs and the Pharmaceutical Sciences, Vol. 126. Marcel Dekker: New York, 
NY. 
3. Charman WN (2000) Lipids, lipophilic drugs and oral delivery-some emerging concepts. 
JPharmSci 89:967-978. 
4. Bowtle W (2000) Lipid formulations for oral drug delivery. Pharm Technol Eur 12(9):20-30. 
5. Attwood D (1994) Microemulsions, in Kreutrer J (ed.) Colloidal Drug Delivery Systems. 
Marcel Dekker: New York, pp. 31-71. 
6. Lawrence MJ (1996) Microemulsions as drug delivery vehicles. Curr Opin Colloid Interface 
Sci 1:826-832. 
7. Pouton CW and Charman WN (1997) The potential of oily formulations for drug delivery 

to the gastrointestinal tract. Adv Drug Del Rev 25:1-2. 
8. Muller RH, Mader K and Gohla S (2000) Solid lipid nanoparticles (SLN) for controlled 
drug delivery. A review of the state of the art. Eur J Pharm Biopharm 50:161-177. 
9. Westesen K (2000) Novel lipid-based colloidal dispersions as potential drug administration 
systems, expectations and reality. Colloid Polym Sci 278:608-618. 
10. Lamprecht A, Bouligand Y and Benoit JP (2002) New lipid nanocapsules exhibit sustained 
release properties for amiodarone. / Control Rel 84:59-68. 
11. Lamprecht A, Saumet JL, Roux J and Benoit JP (2004) Lipid nanocarriers as drug delivery 
systems for Ibuprofen in pain treatment. Intl} Pharm 278(2):407-414. 
12. Razzaque MS, Koji T, Kumatori A and Tagushi T (1999) Cisplatin-induced apoptosis in 
human proximal tubular epithelial cells is associated with the activation of the Fas/Fas 
ligand system. Histochem Cell Biol 111:359-365. 
13. Razzaque MS, Hossain MA, Ahsan N and Tagushi T (2001) Lipid formulations of polyene 
antifungal drugs and attenuation of associated nephrotoxicity. Nephron 89:251-254. 
14. Hossain MA, Maesaki S, Kakeya H, Noda T, Yanagihara K, Sasaki E, Hirakata Y, 
Tomono K, Tashiro T and Kohno S (1998) Efficacy of NS-718, a novel lipid nanosphereencapsulated 
amphotericin B, against cryptococcus neoformans. Antimicrob Agents 
Chemother 42:1722-1725. 
15. Otsubo T, Maesaki S, Yamamoto Y, Tomono K, Tashiro T, Seki J, Tomii Y, Sonoke S and 
Kohno S (1999) In vitro and in vivo activities of NS-718, a new lipid nanosphere incorporating 
amphotericin B, against Aspergillus Fumigatus. Antimicrob Agents Chemother 
43:471-475. 
16. Trotta M, Cavalli R, Carlotti ME, Battaglia L and Debemardi F (2005) Solid lipid microparticles 
carrying Insulin formed by solvent-in-water emulsion-diffusion technique. Int 
J Pharm 288:281-288. 
17. Zarif L (2002) Elongated supramolecular assemblies in drug delivery. / Control Rel Rev 
81:7-23. 
Cochleates as Nanoparticular Drug Carriers 363 
18. Zarif L, Graybill JR, Perlin D and Mannino RJ (2000) Cochleates: New lipid-based drug 
delivery system. / Liposome Res 10(4)523-538. 
19. Papahadjopoulos D, Vail WJ, Jacobson K and Poste G (1975) Cochleate lipid cylinders: 
Formation by fusion of unilamellar lipid vesicles. Biochim Biophys Acta 394(3) :483^91. 
20. Papahadjopoulos D (1978) Large unilamellar vesicles (LUV) and method of preparing 
the same. US Patent 4078052. 
21. Papahadjopoulos D, Nir S and Duzgunes N (1990) Molecular mechanisms of calciuminduced 
membrane fusion. / Bioenerg Biomembr 22(2):157-179. 
22. Wilschut J and Papahadjopoulos D (1979) Ca2+ induced fusion of phospholipid vesicles 
monitored by mixing of aqueous contents. Nature 281(5733):690-692. 
23. Mannino RJ and Gould-Fogerite S (1997) Antigen cochleate formulations for oral and 
systemic vaccination, in Levine MM (ed.) New Generation Vaccines. Marcel Dekker: New 
York, pp. 229-239. 
24. Zarif L and Perlin D (2002) Amphotericin B nanocochleates: From formulation to oral 
efficacy. Drug Del Technol 2(3):34-37. 
25. Zarif L, Graybill JR, Perlin D, Navjar L, Bocanegra R and Mannino RJ (2000) Antifungal 
activity of amphotericin B cochleates against Candida albicans in a mouse model. Antimicrob 
Agents Chemother 44(6):1463-1469. 
26. Walsh TJ, Viviani MA, Arathoon E, Chiou C, Ghannoum M, Groll AH and Odds FC (2000) 
New targets and delivery systems for antifungal therapy. MedMycol 38(Supp l):335-347. 
27. Biodelivery Sciences: www.biodeliverysciences.com 
28. Unger E (2002) Method for delivering bioactive agents using cochleates. US 6403056 Bl. 
29. Zarif L, Jin T, Segarra I and Mannino RJ (2001) New cochleate formulations, process 
of preparation and their use for the delivery of biologically relevant molecules. PCT 
application WO01 /52817 A2. 
30. Zarif L and Mannino RJ (2000) Cochleates: lipid-based vehicles for gene deliveryconcept, 
achievements and future development, in Habib N (ed.) Cancer Gene Therapy: 
Past Achievements and Future Challenges. Kluwer Academic/Plenum Publishers: New 
York, pp. 83-94. 
31. Graham I, Gagne J and Silvius JR (1985) Kinetics and thermodynamics of calciuminduced 
lateral phase separations in phosphatidic acid containing bilayers. Biochemistry 
24(25):7123-7131. 
32. Archibald DD and Mann S (1994) Self-assembled microstructures from 1,2-ethanediol 
suspensions of pure and binary mixtures of neutral and acidic biological galactosylceramides. 
Chem Phys Lipids 69(l):51-64. 
33. Holland JW, Hui C, Cullis PR and Madden TD (1996) Polyethylene glycol)-lipid 
conjugates regulate the calcium-induced fusion of liposomes composed of phosphatidylethanolamine 
and phosphatidylserine. Biochemistry 35(8):2618-2624. 
34. Villardita C, Grioli S, Salmeri G, Nicoletti F and Pennisi G (1987) Multicenter clinical 
trial of brain phosphatidylserine in elderly patients with intellectual deterioration, Clin 
Trials ] 24:84-93. 
35. Crook TH, Tinklenberg J, Yesavage J, Petrie W, Nunzi MG and Massari DC (1991) Effects 
of phosphatidylserine in age-associated memory impairment. Neurology 41:644-649. 
364 Zarif 
36. Engle RR, Satzger W and Gunther W (1992) Double-blind cross-over study of phosphatidylserine 
vs. placebo in patients with early dementia of the Alzheimer type. Eur 
Neuropsychopharmacol 2:149-155. 
37. Amaducci L (1988) Phosphatidylserine in the treatment of Alzheimer's disease: Results 
of a multicenter study. Psychopharmacol Bull 24(1):130-134. 
38. Cennachi T, Bertoldin T, Farina C, Fiori MG and Crepaldi G (1993) Cognitive decline in 
the elderly: A double blind, placebo-controlled multicenter study on efficacy of phosphatidylserine 
administration. Aging 5(2):123-133. 
39. Guidin J et al. (1995) Effect of soy lecithin phosphatidylserine (PS) complex on memory 
impairment and mood in the functioning elderly. Dept Geriatrics, Kaplan Hospital, 
Rehovot, Israel. 
40. Maggioni M, Picotti GB, Bondiolotti GP et al. (1990) Effects of phosphatidylserine therapy 
in geriatric patients with depressive disorders. Acta Psychiatr Scand 81:265-270. 
41. Nerozzi D, Aceti F, Melia E, Magnani A, Marino R, Genovesi G, Amalfitano M, 
Cozza G, Murgiano S, De Giorgis G, et al. (1987) Phosphatidylserine and Memory 
Disorders in the Aged. Clin Ter 120(5):399^04. 
42. Monteleone P, Beinat L, Tanzillo C, Maj M and Kemali D (1990) Effects of phosphatidylserine 
on the neuroendocrine response to physical stress in humans. Neuroendocrinology 
52:243-248. 
43. Monteleone P et al. (1992) Blunting by chronic phosphatidylserine administration of the 
stress-induced activation of the hypothalamo-pituitary-adrenal axis in healthy men. Eur 
] Clin Pharmacol 41:385-388. 
44. Funfgeld EW, Baggen M, Nedwidek P, Richstein B and Mistlberger G (1989) Doubleblind 
study with phosphatidylserine (PS) in parkinsonian patients with senile dementia 
of Alzheimer's type (SDAT). Prog Clin Res 317:1235-1246. 
45. Zarif L and Tan F (2003) Cochleates made with purified soy phosphatidylserine. 
US2003 / 0219473 Al. 
46. Duzgunes N, Nir S, Wischut J, Bentz J, Newton C, Portis A and Papahadjopoulos 
D (1981) Calcium- and magnesium induced fusion of mixed phosphatidylserine/ 
phosphatidylcholine vesicles: Effect of ion-binding, f Membr Biol 59:115-125. 
47. Portis A, Newton C, Pangborn W and Papahadjopoulos D (1979). Studies on the mechanism 
of membrane fusion: Evidence for an intermembrane Ca2+-phospholipid complex, 
synergism with Mg2+, and inhibition by spectrin. Biochemistry 18:780-790. 
48. Santangelo R, Paderu P, Delmas G, Chen ZW Mannino R, Zarif L and Perlin D (2000) 
Efficacy of oral cochleates amphotericin b in a mouse model of systemic candidiasis. 
Antimicrob Agents Chemother 44(9):2356-2360. 
49. Zarif L, Segarra I, Jin T, Scolpino A, Hyra D, Daublin P, Krause S, Perlin DS, Lambros C, 
Graybill JR and Mannino RJ (1999) Lipid-based cochleate system for oral and systemic 
delivery of drugs. AAPS Eastern Regional Meeting and Exposition. 
50. Segarra I, Hyra-Movshin DA, Chen ZW, Santangelo R, Perlin D, Paderu P, Mannino RJ 
and Zarif L (2000) AmB Cochleates, a new lipid-based formulation for amphotericin 
Cochleates as Nanoparticular Drug Carriers 365 
B: From IV pharmacokinetics to oral efficacy. Millenial World Congress of Pharmaceutical 
Sciences, San Franscisco, CA, April, pp. 124. 
51. Segarra I, Jin T, Hyra D, Mannino RJ and Zarif L (1999) Oral administration of amphotericin 
B with a new AmB-cochleate formulation: Tissue distribution after single and 
multiple oral dose. 1CAAC 39:Abs 1940. 
52. Segarra I, Movshin D, Mannino RJ and Zarif L (2000) Pharmacokinetics and tissue distribution 
of amphotericin B in Mice after oral administration of AmB cochleates, a new 
effective lipid-based formulation for the oral treatment of systemic fungal infections. 
ICAAC 40:Abs 861. 
53. Segarra I, Chen ZW, Movshin DA, Tan F, Mannino RJ and Zarif L (2000) Tissue distribution 
of oral amphotericin B lipid-based cochleate formulation: Comparison with 
AmBisome. 27th International Symposium on Controlled Release ofBioactive Materials, Paris 
France, pp. 67-68. 
54. Zarif L, Segarra I, Jin T, Hyra D and Mannino RJ (1999) Amphotericin B cochleates as 
a novel oral delivery system for the treatment of fungal infections. 26th International 
Symposium on Controlled Release ofBioactive Materials. Boston, MA, June 20-23. 
55. Perlin D, Santangelo R, Mannino R and Zarif L (2000) Oral delivery of cochleates containing 
amphotericin B (CAMB) is highly effective in a candidiasis murine model, Focus 
Fungal Infect. 
56. Zarif L, Segarra I, Jin T, Hyra D, Perlin D, Graybill JR and Mannino JR (1999) Oral and 
systemic delivery of amphotericin B mediated by cochleates. AAPS Annual Meeting and 
Exposition, November. 
57. Zarif L, Jin T, Scolpino A and Mannino RJ (1999) Are cochleates the new lipid-based 
carrier for oral drug delivery? 39th ICAAC, San Francisco, CA, September 26-29. 
58. Delmas G, Perlin D, Chen ZW and Zarif L (2001) Amphotericin B cochleates: Evaluation 
for the oral treatment of aspergillosis in murine model, The 28th International Symposium 
of Controlled Release of Bioactive Materials, San Diego, CA, June 23-29, pp. 433^34. 
59. Delmas G, Park S, Chen ZW, Tan F, Kashiwazaki R, Zarif L and Perlin DS (2002) Efficacy of 
orally delivered cochleates containing amphotericin B in a murine model of aspergillosis. 
Antimicrob Agents Chemother 46(8):2704-2707. 
60. Graybill JR, Navjar L, Bocanegra R, Scolpino A, Mannino RJ and Zarif L (2000) A new 
lipid vehicle for amphotericin B, Abstract, 39th ICAAC, San Franscisco, CA, September, 
Abs 583. 
61. Delmarre D, Lu R, Taton N, Krause-Elsmore S, Gould-Fogerite S and Mannino RJ (2004) 
Cochleate-mediated delivery: Formulation of hydrophobic drugs into cochleate delivery 
vehicles: A simplified protocol & bioral formulation kit. Drug Del Techno 4(l):64-69. 
62. Ramani K and Balasubramanian S (2003) Fluorescence properties of Laurdan in cochleate 
phases. Biochim Biophys Acta 1618(l):67-78. 
63. Rex JH, Walsh TJ, Sobel JD, Filler SG, Pappas PG, Dismukes WE and Edwards JE (2000) 
Practice guidelines for the management of candidiasis. Infectious Diseases Society of 
America. Clin Infect Dis 30(4):662-678. 
64. Saag MS, Graybill RJ, Larsen RA, Pappas PG, Perfect JR, Powderly WG, Sobel JD and 
Dismukes WE (2000) Practice guidelines for the management of cryptococcal disease. 
Infectious Diseases Society of America. Clin Infect Dis 30(4):710-718. 
366 Zarif 
65. Stevens DA, Kan VL, Judson MA, Morrison VA, Dummer S, Dening DW, Bennett JE, 
Walsh TJ, Patterson TF and Pankay GA (2000) Practice guidelines for diseases caused by 
Aspergillus. Infectious Diseases Society of America. Clin Infect Dis 30(4):696-709. 
66. Hiemenz JW and Walsh TJ (1996) Lipid formulations of amphotericin B: Recent progress 
and future directions. Clin Infect Dis 22(Suppl 2):133-144. 
67. Graybill JR, Najvar LK, Bocanegra R, Scolpino A, Mannino RJ and Zarif L (1999). 
Cochleate: A new lipid vehicle for amphotericin B. ICAAC 39:Abs 2009. 
68. Zarif L, Graybill J, Najvar L, Perlin D and Mannino RJ (2000) Amphotericin B cochleates: 
Novel lipid-based drug delivery system for the treatment of systemic fungal infections., 
14th ISHAM World Congress, May 8-12, Buenos Aires, Argentina. 
69. Segarra I, Movshin DA and Zarif L (2002) Extensive tissue distribution of amphotericin 
B after intravenous administration in cochleate vehicle to mice. 29th International Symposium 
on Controlled Release of Bioactive Materials, Seoul, Korea. 
70. Segarra I, Movshin D and Zarif L (2002) Pharmacokinetics and tissue distribution after 
intravenous administration of a single dose of amphotericin B cochleates, a new lipidbased 
delivery system. / Pharm Sci 91(8):1827-1837. 
71. Legrand P, Vertut-Doi A and Bolard J (1996) Comparative internalization and recycling 
of different amphotericin B formulations by a macrophage-like cell line. / Antimicrob 
Chemother 37:519-533. 
72. Bratosin D, Mazurier J, Tissier JP, Slomianny C, Estaquier J, Russo-Marie F, Huart JJ, 
Freyssinet JM, Aminoff D, Ameisen JC and Montreuil J (1997) Molecular mechanism of 
erythrophagocytosis. Characterization of the senescent erythrocytes that are phagocytized 
by macrophages. CR Acad Sci Paris Sciences de la Vie/Life Sci 320:811-818. 
73. Popescu C, Adams L, Franzblau S and Zarif L (2001) Cochleates potentiate the efficacy 
of the antimycobacterial drug, clofazimine. ICAAC 41:Abs 2278. 
74. Jin T (2003) Cochleates without metal cations as bridging agents. US Patent application 
10/636,522. 
75. Slayton W, Anstine D, Lakhdir F, Sleasman J and Neiberger R (1996) Tetany in a child 
with AIDS receiving intravenous tobramycin. South Med J 89:1108-1110. 
76. Keating MJ, Sethi MR, Bodey GP and Samaan NA (1977) Hypocalcemia with hypopara 
thyroidism and renal tubular dysfunction associated with aminoglycoside therapy. 
Cancer 39:1410-1414. 
77. RRC (1990) New (Ed.), Liposomes, a practical approach, IRL Press, Oxford University 
Press, New York. 
78. Gould-Fogerite S, Mazurkiewicz JE, Raska K Jr, Voelkerding K, Lehman JM and Mannino 
RJ (1989) Gene 84(2):429-438. 
79. Perez O, Brach G, Lastre M, Mora N, Del Campo J, Gil D, Zayas C, Acevedo R, 
Gonzales D, Lopez J, Taboada C and Solis RL (2004) Novel adjuvant based on 
a proteoliposome-derived cochleate structure containing native polysaccharide as a 
pathogen-associated molecular pattern. Immunol Cell Biol 82(6):603-610. 
17 
Aerosols as Drug Carriers 
N. Renee Labiris, Andrew P. Bosco 
and My ma B. Dolovich 
1. Introduction 
As the end organ for the treatment of local diseases or as the route of administration 
for systemic therapies, the lung is a very attractive target for drug delivery (Table 1). 
The lung provides direct access to the site of disease for the treatment of respiratory 
illness, without the inefficiencies and unwanted effects of systemic drug delivery. 
In addition, it provides an enormous surface area and a relatively low enzymatic 
environment for the absorption of drugs to treat systemic diseases (Table 1). 
Inhaled medications have been available for many years for the treatment of 
lung diseases. Inhalational delivery has been widely accepted as being the optimal 
route of administration of first line therapy for asthmatic and chronic obstructive 
pulmonary diseases. Drug formulation plays an important role in producing an 
effective inhalable medication. In addition to being pharmacologically active, it is 
important that a drug be efficiently delivered into the lungs, to the appropriate site 
of action and remain in the lungs until the desired pharmacological effect occurs. 
A drug designed to treat a systemic disease, such as insulin for diabetes, must be 
deposited in the lung periphery to ensure maximum systemic bioavailability. For 
gene therapy, anti cancer or anti infective treatment, cellular uptake and prolonged 
residence in the lungs of the drug may be required to obtain the optimal therapeutic 
effect. Thus, a formulation that is retained in the lungs for the desired length of time 
and avoids the clearance mechanisms of the lung may be necessary. 
The human lung contains airways and approximately 300 million alveoli with 
a surface area of 140 m2, equivalent to that of a tennis court.1 As a major port of 
367 
368 Labiris, Bosco & Dolovich 
Table 1 
disease. 
Advantages of pulmonary delivery of drugs to treat respiratory and systemic 
Treatment of respiratory diseases Treatment of systemic diseases 
Deliver high drug concentrations directly 
to the disease site 
Minimizes risk of systemic side effects 
Rapid clinical response 
Bypass the barriers to therapeutic 
efficacy, such as poor gastrointestinal 
absorption and first-pass metabolism in 
the liver 
Achieve a similar or superior therapeutic 
effect at a fraction of the systemic dose. 
For example, oral salbutamol 2-4 mg is 
therapeutically equivalent to 100-200 /xg 
byMDI 
A non-invasive Needle-free delivery 
system. 
Suitable for a wide range of substances 
from small molecules to very large 
proteins 
Enormous absorptive surface area 
(140 m2) and a highly permeable 
membrane (0.2 to 0.7 /xm thickness) in 
the alveolar region. 
Large molecules with very low 
absorption rates can be absorbed in 
significant quantities; the slow 
mucociliary clearance in the lung 
periphery results in prolonged residency 
in the lung. 
A less harsh, low enzymatic 
environment 
Avoids first-pass metabolism. 
Reproducible absorption kinetics. 
Pulmonary delivery is independent 
of dietary complications, extracellular 
enzymes and inter-patient metabolic 
differences that affect gastrointestinal 
absorption. 
entry, the lung has evolved to prevent the invasion of unwanted airborne particles 
from entering into the body. Airway geometry, humidity, mucociliary clearance 
and alveolar macrophages play a vital role in maintaining the sterility of the lung, 
and consequently, they can be barriers to the therapeutic effectiveness of inhaled 
medications. 
The size of the drug particle can play an important role in avoiding the physiological 
barriers of the lung and targeting to the appropriate lung region (Fig. 1). 
Nanoparticles are solid colloidal particles ranging in size from 10 to 1000 nm.2 
Studies have demonstrated that they are taken up by macrophages, cancer cells, 
and epithelial cells.3-6 Their small size ensures the particles containing the active 
pharmacological ingredient will reach the alveolar regions. However, the use of an 
aerosol delivery system that generates nano-sized particles for inhalation, places 
these particles at risk of being exhaled, leaving very few drug particles to be 
deposited in the periphery of the lung. Residence time is not long enough for the 
particles to be deposited by sedimentation or diffusion.7 
DIFFUSION 
Aerosols as Drug Carriers 369 
SEDIMENTATION INITIAL IMPACTION 
05 ID 2.0 5.0 
AERODYNAMIC DIAMETER pm (Mkrara) 
Fig. 1. Relationship between particle size and lung deposition. 
SOB 
105 
2. Pulmonary Drug Delivery Devices 
The origin of inhaled therapies can be traced back 4000 years ago to 
India, where people smoked the leaves of the Atropa belladonna plant to 
suppress cough. In the 19th and early 20th centuries, asthmatics smoked 
asthma cigarettes that contained stramonium powder mixed with tobacco 
to treat the symptoms of their disease. Modern inhalation devices can be 
divided into three different categories (Fig. 2), the refinement categories 
(Fig. 2), the refinement of the nebulizer and the of compact portable 
devices, the pressurized metered dose inhaler (pMDI), and the dry powder 
inhaler (DPI). The advantages and disadvantages of each are summarized in 
Table 2. 
2.1. Nebulizers 
Nebulizers have been used for many years to treat asthma and other respiratory 
diseases. There are 2 basic types of nebulizers, jet and ultrasonic nebulizers. The 
jet nebulizer functions by the Bernoulli principle by which compressed gas (air 
or oxygen) passes through a narrow orifice, creating an area of low pressure at 
the outlet of the adjacent liquid feed tube. This results in the drug solution being 
drawn up from the fluid reservoir and shatter into droplets in the gas stream. The 
ultrasonic nebulizer uses a piezoelectric crystal, vibrating at a high frequency (usually 
1 to 3 MHz), to generate a fountain of liquid in the nebulizer chamber; the 
higher the frequency, the smaller the droplets produced. Nebulizers can aerosolize 
3 70 Labiris, Bosco & Dolovich 
Glass Nebulizer 
(Late 19* century) 
Hand Bulb Nebulizer 
(1938) 
Adaptive 
Aerosol 
Delivery 
Metered Dose Inhalers (MDI) 
(1956, CFC prcmellant) 
Metered Dose 
Liquid Inhalers 
Breath-Actuated 
MDI 
Add-On 
Devices 
CFC-Free 
MDI 
Dry Powder Inhaler 
(DPI) 
Passive Active 
Fig. 2. Evolution of pulmonary delivery devices. 
most drug solutions and provide large doses with very little patient coordination 
or skill. However, treatments using these nebulizers can be time consuming and 
inefficient, with large amounts of drug wastage e.g. 50% loss with continuously 
operated nebulizers.8 Most of the prescribed drug never reaches the lung with nebulization. 
The majority of the drug is either retained within the nebulizer (referred 
to as residual or dead volume) or released into the environment during expiration. 
On average, only 10% of the dose placed in a continuous output jet nebulizer is 
actually deposited in the lungs.8 Advances in technology have led to the development 
of novel nebulizers that reduce drug wastage and improve delivery efficiency. 
Breath-enhanced jet nebulizers such as the Pari LC Star, (PARI, Germany) increase 
aerosol output by directing auxiliary air, entrained during inspiration, through 
the nebulizer, causing more of the generated aerosol to be swept out of the nebulizer 
and available for inhalation. Drug wastage during exhalation is reduced to 
the amount of aerosol produced by the jet airflow rate that exceeds the storage 
volume of the nebulizer. Adaptive aerosol delivery (Halolite, Medic-Aid, Bognor 
Regis, UK) monitors a patient's breathing pattern in the first 3 breaths and then targets 
the aerosol delivery into the first 50% of each inhalation. This ensures that the 
aerosol is delivered to the patient during inspiration only, thereby eliminating drug 
loss during expiration that occurs with continuous output nebulizers.9 A number 
of metered dose liquid inhalers, including AERx (Aradigm, Hayward, CA), Aero- 
Dose (AeroGen, Sunnyvale, CA) and Respimat (Boehringer Ingelheim, Ingelheim 
Rhein, Germany), have been developed to produce a fine aerosol in the respirable 
Aerosols as Drug Carriers 371 
Table 2 Advantages and disadvantages of inhalation devices. 
Inhalation device Advantages Disadvantages 
Nebulizers (jet, ultrasonic) no specific inhalation 
technique or coordination 
required 
aerosolizes most drug 
solutions 
delivers large doses 
suitable for infants and 
people too sick or 
physically unable to use 
other devices 
time consuming 
bulky 
non-portable 
contents easily 
contaminated 
relatively expensive 
poor delivery efficiency 
drug wastage 
wide performance 
variation between 
models and operating 
conditions 
pressurized Metered Dose 
Inhalers (pMDI) 
Dry Powder Inhalers (DPI) 
compact 
portable 
multi-dose (-200 
doses) 
inexpensive 
sealed environment (no 
degradation of drug) 
reproducible dosing 
compact 
portable 
breath actuated 
easy to use 
no hand-mouth 
coordination required 
inhalation technique 
and patient coordination 
required 
high oral deposition 
maximum dose of 5 mg 
limited range of drugs 
available 
respirable dose 
dependent on IFR* 
humidity may cause 
powders to aggregate 
and capsules to soften 
dose lost if patient 
inadvertently exhales 
into the DPI 
most DPIs contain 
lactose 
*IFR = Inspiratory Flow Rate 
range by forcing the drug solution through an array of nozzles, using vibrating 
mesh or electronic micropump platforms with 30 to 75% of the emitted dose being 
deposited in the lungs.10,11 
2.2. Metered-dose inhalers 
The pressurized metered-dose inhaler (pMDI) was a revolutionary invention that 
overcame the problems of the hand-bulb nebulizer, and it is the most widely 
used aerosol delivery device today. The pMDI emits a drug aerosol driven by 
372 Labiris, Bosco & Dolovich 
propellants, such as chlorofluorocarbons (CFC) and more recently, hydrofluoroalkanes 
(HFAs) through a nozzle at high velocity (>30m/sec). pMDIs deliver only a 
small fraction of the drug dose to the lung. Typically, only 10 to 20% of the emitted 
dose is deposited in the lung.12 The high velocity and large particle size of 
the spray causes approximately 50% to 80% of the drug aerosol to impact in the 
oropharygeal region.13 Hand-mouth discoordination is another obstacle in the optimal 
use of the pMDI. Crompton and colleagues14 found 51% of patients experienced 
problems coordinating the actuation of the device with inhalation, 24% of 
patients halted inspiration upon firing the aerosol into the mouth, and 12% inspired 
through the nose instead of the mouth when the aerosol was actuated into the 
mouth. 
The delivery efficiency of a pMDI depends on a patient's breathing pattern, 
inspiratory flow rate and hand-mouth coordination. The studies by Bennett15 and 
Dolovich16 demonstrated that for any particle size between 1 to 5 /tm mass median 
aerodynamic diameter (MMAD), deposition was more dependent on inspiratory 
flow rate than any other variable. Fast inhalations (>60 L/min) result in a reduced 
peripheral deposition because the aerosol is more readily deposited by inertial 
impaction in the conducting airway and oropharyngeal regions. When aerosols 
are inhaled slowly, deposition by gravitational sedimentation in peripheral lung 
regions are enhanced.17 Peripheral deposition has also been shown to increase with 
an increase in tidal volume and a decrease in respiratory frequency. As the inhaled 
volume is increased, aerosols are able to penetrate more distally into the lungs.18 
A period of breath holding on completion of inhalation enhances deposition of 
particles in the periphery, thus preventing the particles from being exhaled during 
the expiratory phase. Thus, the optimal conditions for inhaling pMDI aerosols are 
from a starting volume equivalent to the functional residual capacity, the actuation 
of the device at the start of inhalation, inspiratory flow rate of <60 L/min, followed 
by a 10 second breath-hold at the end of inspiration.17,19 
Spacer tubes, valved holding chambers and mouthpiece extensions have been 
developed to eliminate coordination requirements and reduce the amount of drug 
deposited in the oropharynx, by decreasing the particle size distribution and slowing 
the aerosol's velocity. Spacer geometry and materials of manufacture influence 
the quality and quantity of aerosol available. The aerosols from a pMDI and the 
holding chamber are finer than that with the pMDI alone, with an approximate 
25% decrease in the mass median aerodynamic diameter (MMAD), compared with 
the original aerosol.20,21 This finer aerosol is more uniformly distributed in the normal 
lung, with increased delivery to the peripheral airway. However, in patients 
with airway obstructions, the addition of a holding chamber to the pMDI may not 
change the distribution of the aerosol.22 
Aerosols as Drug Carriers 373 
2.3. Dry powder inhalers 
Dry powder inhalers (DPIs) were designed to eliminate the coordination difficulties 
associated with the pMDI. There are a wide range of DPI devices on the market from 
single-dose devices loaded by the patient (e.g. Aerolizer from Novartis, Rotahaler 
from GSK, Ware UK) to multi unit dose devices provided in a blister pack (e.g. 
Diskhaler, GSK, Ware UK), multiple unit doses sealed in blisters on a strip that 
moves through the inhaler (e.g. Diskus, GSK, Ware UK) or reservoir-type (bulk 
powder) systems (e.g. Turbuhaler, AstraZeneca, Lund Sweden). 
Lung deposition varies among the different DPIs. Approximately 12% to 40% 
of the emitted dose is delivered to the lungs with 20 to 25% of the drug being 
retained within the device.10,23,24 Poor drug deposition with DPIs can be attributed 
to inefficient deaggregation of the fine drug particles from coarser carrier lactose 
particles or drug pellets. Slow inspiratory flow rate, high humidity and rapid, large 
changes in temperature are known to affect drug deaggregation and hence the efficiency 
of pulmonary drug delivery with DPIs.25,26 With most DPIs, drug delivery 
to the lungs is augmented by fast inhalation. Borgstrom and colleagues27 demonstrated 
that increasing inspiratory flow from 35L/min to 60L/min through the 
Turbuhaler7, increased the total lung dose of terbutaline from 14.8% of nominal 
dose to 27.7%. This is in contrast to the MDI which requires slow inhalation and 
breath holding to enhance lung deposition of the drug. Each DPI has a different 
air flow resistance that governs the required inspiratory effort.28,29 The higher the 
resistance of the device, the more difficult it is to generate an inspiratory flow great 
enough to achieve the maximum dose from the inhaler.30-32 However, deposition 
in the lung tends to increase when using high resistance inhalers.32-36 
Active DPIs are being investigated to reduce the importance of a patient's inspiratory 
effort. By adding either a battery driven propeller that aids in the dispersion 
of the powder (Spiros, Elan Pharmaceuticals, San Diego, CA), or using compressed 
air to aerosolize the powder and converting it into a standing cloud in a holding 
chamber, the generation of a respirable aerosol becomes independent of a patient's 
inspiratory effort (Inhance Pulmonary Delivery System, Nektar Therapeutic, San 
Carlos, CA). 
3. Aerosol Particle Size 
Aerosol particle size is one of the most important variables in defining the dose 
deposited and the distribution of drug aerosol in the lung (Fig. 3). Fine aerosols 
are distributed on peripheral airways, but deposit less drug per unit surface area 
than larger particle aerosols which deposit more drug per unit surface area, but on 
3 74 Labiris, Bosco & Dolovich 
(a) JOO*. 
FREQUENCY DISTRIBUTION 
% 50 _ 
NUMBERVOLUME 
{MASS] 
0.1 i )0 100 
AERODYNAMfC DIAMETER JJm 
(b) CUMULATIVE DISTRIBUTION 
100 | - / > 
% 50 
NUMBER-^ /--VOLUME CMAS5) 
/ MMAD = 2.25 ftm 
_L 
0.1 i 10 
AERODYNAl-flC DIAMETER Jtm 
100 
Fig. 3. Frequency (a) and cumulative (b) distribution curves for Beclovent MDI used with 
an Aerochamber, in terms of number of particles and volume (mass) of particles vs. particle 
aerodynamic diameter. The volume distribution curves are displaced to the right of the 
number distribution curves. The smaller number of large particles within the aerosol carry the 
greater mass of the drug; this is reflected in the larger, second peak of the volume distribution 
curve, which corresponds to the smaller second peak of the number distribution curve. 
MMAD is read from the cumulative distribution curve at the 50% point and if the distribution 
is log-normal, the GSD can be calculated as the ration of the diameter at the 84.1% point to 
the MMAD. Particle distribution was measured using the Anderson Cascade Impactor.105 
the larger, more central airways.37 Most therapeutic aerosols are nearly always heterodisperse, 
consisting of a wide range of particle sizes. These aerosols are described 
by the log-normal distribution, with the log of the particle diameters plotted against 
particle number, surface area or volume (mass) on a linear or probability scale and 
expressed as absolute values or cumulative %. Since delivered dose is very important 
when studying medical aerosols, particle number may be misleading as smaller 
particles contain less drug than larger ones. Particle size is defined from this distribution 
by several parameters. Mass median diameter of an aerosol refers to the 
Aerosols as Drug Carriers 375 
particle diameter that has 50% of the aerosol mass residing above and 50% of its 
mass below it. The aerodynamic diameter relates the particle to the diameter of a 
sphere of unit density that has the same settling velocity as the particle of interest, 
regardless of its shape or density. MMAD is read from the cumulative distribution 
curve at the 50% point (Fig. 3). Geometric standard deviation (GSD) is a measure of 
the variability of the particle diameters within the aerosol, and is calculated from 
the ratio of the particle diameter at the 84.1% point on the cumulative distribution 
curve to the MMAD. For a log-normal distribution, the GSD is the same for the 
number, surface area or mass distributions. A GSD of 1 indicates a monodispersed 
aerosol, while a GSD of > 1.2 indicates a heterodispersed aerosol. 
Particles can be deposited by inertial impaction, gravitational sedimentation 
or diffusion (Brownian motion), depending on their size. While deposition occurs 
throughout the airways, inertial impaction usually occurs in the first 10 generations 
of the lung, where air velocity is high and airflow is turbulent.38 Most particles above 
10 /xm are deposited in the oropharyngeal region with a large amount impacting 
on the larynx, particularly when the drug is inhaled from devices requiring a high 
inspiratory flow rate (DPIs) or when the drug is dispensed from a device at a high 
forward velocity (MDIs).39,40 The large particles are subsequently swallowed and 
contributed minimally, if at all, to the therapeutic response. In the tracheobronchial 
region, inertial impaction also plays a significant role in the deposition of particles, 
particularly at bends and airway bifurcations. Deposition by gravitational sedimentation 
predominates in the last 5 to 6 generation of airways (smaller bronchi 
and bronchioles), where air velocity is low.38 In the alveolar region, air velocity 
is negligible and thus the contribution to deposition by inertial impaction is also 
negligible. Particles in this region have a longer residence time and are deposited 
by both sedimentation and diffusion. Particles not deposited during inhalation are 
exhaled. Deposition due to sedimentation affects particles down to 0.5 ^tm in diameter, 
whereas below 0.5 /xm, the main mechanism for deposition is by diffusion. 
Targeting the aerosol to conducting or peripheral airways can be accomplished 
by altering the particle size of the aerosol. It is difficult to predict the actual site 
of deposition, since airway calibre and anatomy differ among people. However, 
in general, aerosols with a MMAD of 5 to 10 /xm are mainly deposited in the large 
conducting airways and the oropharyngeal region.41 Particles 1 to 5 /xm in diameter 
are deposited in the small airways and alveoli with greater than 50% of the 3 /tm 
diameter particles being deposited in the alveolar region. In the case of pulmonary 
drug delivery for systemic absorption, aerosols with a small particle size would 
be required to ensure peripheral penetration of the drug.42 Particles <3 /xm have 
approximately 80% chance of reaching the lower airways, with 50 to 60% being 
deposited in the alveoli.43'44 Nanoparticles <100nm are deposited mainly in the 
alveolar region. 
376 Labiris, Bosco & Dolovich 
4. Targeting Drug Delivery in the Lung 
The therapeutic effect of aerosolized therapies is dependent on the dose deposited 
and its distribution within the lung. If a drug aerosol is delivered at a suboptimal 
dose or to a part of the lung, devoid of the targeted disease or receptors, the 
effectiveness of therapy may be compromised. For example, the receptors for the fc 
agonist, salbutamol and the muscarine (M3) agonist, ipratropium bromide, are not 
uniformly distributed throughout the lung. Autoradiographic studies have shown 
P2 adrenergic receptors are present in high density in the airway epithelium from 
the large bronchi to the terminal bronchioles. Airway smooth muscle has a lower 
/S-receptor density, greater in the bronchioles than bronchi.45 However, greater than 
90% of all /3 receptors are located in the alveolar wall, a region where no smooth 
muscle exists and whose functional significance is unknown. Another autoradiographic 
study has shown a high density of M3 receptors in submucosal glands 
and airway ganglia, and a moderate density in smooth muscles throughout the 
airways, nerves in intrapulmonary bronchi and in alveolar walls.46 The location of 
these receptors in the lung suggests that ipratropium bromide needs to be delivered 
to the conducting airways, while salbutamol requires a more peripheral delivery 
to the medium and small airways to produce a therapeutic effect. 
Since particle size affects the lung deposition of an aerosol, it can also influence 
the clinical effectiveness of a drug. Rees et al. reported the varying clinical effect 
of 250 /xg of aerosolized terbutaline from a pMDI, given in three different particle 
sizes of <5 /xm, 5 to 10 /xm, and 10 to 15 /xm.47 In asthmatics, the greatest increase in 
forced expiratory volume in one second (FEVi) was found with the smallest particle 
size (<5/xm), suggesting that the smaller particle aerosol was considerably more 
effective than larger particle size aerosols in producing bronchodilation, since it has 
the best penetration and retention in the lungs in the presence of airway narrowing. 
Using three monodisperse salbutamol aerosols (MMAD of 1.5 /xm, 2.8 /xm, 5 Aim), 
Zanen and colleagues demonstrated in patients with mild to moderate asthma 
that the 2.8 /xm particle size aerosol produced a superior bronchodilation, compared 
with the other two aerosols.48 In patients with severe airflow obstruction 
(FEVi < 40%), Zanen et al. demonstrated that the optimal particle size for /J2 agonist 
or anticholinergic aerosols is approximately 3 /xm.49 They examined the effect 
on lung function of equal doses of three different sizes of monodisperse aerosols, 
1.5 /xm, 2.8 /xm and 5 /xm, of salbutamol and ipratropium bromide. Their findings 
suggest that small particles penetrate more deeply into the lung and more effectively 
dilate the small airways than larger particles, which are filtered out in the 
upper airways. The 1.5 /xm aerosol induced significantly less bronchodilation than 
the 2.8 /xm aerosol, suggesting that this fine aerosol may be deposited too peripherally 
to be effective, since smooth muscle is not present in the alveolar region. 
Aerosols as Drug Carriers 377 
The optimal site of deposition in the respiratory tract for aerosolized antibiotics 
depends on the infection being treated. Pneumonias represent a mixture of purulent 
tracheobronchitis and alveolar infection. Successful therapy would theoretically 
require the antibiotic to be evenly distributed throughout the lungs. However, those 
confined to the alveolar region would most likely benefit from a greater peripheral 
deposition. Pneumocystis carinii pneumonia, the most common life-threatening 
infection among patients infected with HIV, is found predominately within the 
alveolar spaces, with relapses occurring in the apical region of the lung after treatment 
with inhaled pentamidine given as a 1 fim MMAD aerosol.50 The mechanism 
suggested for this atypical relapse is the poorer apical deposition of the aerosol. 
Regional changes in intrapleural pressure result in the lower lung regions receiving 
relatively more of the inspired volume than the upper lung, when sitting in an 
upright position or standing. This influence on deposition has been shown to occur 
in an experimental lung model, analyzing sites of aerosol deposition in a normal 
lung. The experiment showed a 2:1 ratio in the overall deposition for a 4 /xm aerodynamic 
diameter aerosol between the lower and upper lobes when in the upright 
position.51 
Chronic lung infection with Pseudomonas aeruginosa, in patients with cystic 
fibrosis or non-CF bronchiectasis, resides in the airway lumen with limited invasion 
of the lung parenchyma.52'53 Infection starts in the smaller airways, the bronchioles, 
and moves into the larger airways. The optimal site of deposition for inhaled 
antimicrobial therapy would, therefore, be a uniform distribution on the conducting 
airways. Mucus plugs in the bronchi and bronchioles may prevent deposition 
of even small particle aerosols in regions distal to the airway obstruction, possibly 
the regions of highest infection, and thereby limiting the therapeutic effectiveness 
of the aerosolized antibiotic.54-56 
Until recently, aerosol drug delivery has been limited to topical therapy for 
the lung and nose. The major contributing factor to this restriction was the inefficiencies 
of available inhalation devices that deposit only 10% to 15% of the emitted 
dose in the lungs. While appropriate lung doses of steroids and bronchodilators can 
be achieved with these devices, for systemic therapies, large amounts of the drug 
are necessary to achieve therapeutic drug levels systemically. Recent advances in 
aerosol and formulation technologies have led to the development of delivery systems 
that are more efficient and that which produce small particle aerosols, allowing 
higher drug doses to be deposited in the alveolar region of the lungs, where they 
are available for systemic absorption. 
Most macromolecules cannot be administered orally because proteins are 
digested before they are absorbed into the bloodstream. In addition, their large 
size prevents them from naturally passing through the skin or nasal membrane; 
therefore, they cannot be administered intranasally or transdermally without the 
378 Labiris, Bosco & Dolovich 
use of penetration enhancers. Thus, the easiest route of administration for proteins 
has been through intravenous or intramuscular/subcutaneous injection. It has been 
known for many years that proteins can be absorbed from the lung as demonstrated 
with insulin in 1925.57 Macromolecules < 40 kiloDaltons (kDa) (<5-6nm 
in diameter) appear rapidly in the blood following inhalation into the airways. 
Insulin which has a molecular weight (mw) of 5.7 kDa and a diameter of 2.2 nm 
peaks in the blood 15 to 60 min after inhalation.58-62 Macromolecules >40 kDa (>5- 
6 nm in diameter) are slowly absorbed over many hours; inhaled albumin (68 kDa) 
and alphai-antitrypsin (45-51 kDa) have a Tmax of 20hrs and between 12 to 48hrs 
respectively.63 
The lung is the only organ through which the entire cardiac output passes. 
Before the inhaled drug can be absorbed into the blood from the lung periphery, 
it has several barriers to overcome such as lung surfactant, surface lining fluid, 
epithelium, interstitium and basement membrane, and the endothelium. Drug 
absorption in the lung periphery is regulated by a thin alveolar-vascular permeable 
barrier. An enormous alveolar surface area with epithelium, consisting of a 
thin single cellular layer (0.2 to 0.7 /xm thickness), promotes efficient gas exchange 
through passive transport, but also provides a mechanism for efficient drug delivery 
into the bloodstream.64 Although the mechanism of absorption is unknown, 
it has been hypothesized that macromolecules either pass through the cells via 
absorptive transcytosis (adsorptive or receptor mediated), paracellular transport 
between bijunctions or trijunctions or through large transitory pores in the epithelium 
caused by cell injury or apoptosis.65 Thus, the high bioavailability of macromolecules 
deposited in the lung (10 to 200 times greater than nasal and gastrointestinal 
values) may be due to its enormous surface area, very thin diffusion layer, 
slow surface clearance and anti-protease defense system. 
5. Clearance of Particles from the Lung 
Like all major points of contact with the external environment, the lung has evolved 
to prevent the invasion of unwanted airborne particles from entering into the body. 
Airway geometry, humidity and clearance mechanisms contribute to this filtration 
process. The challenge in developing therapeutic aerosols is to produce an aerosol 
that eludes the lung's various lines of defense. 
5.1. Airway geometry and humidity 
Progressive branching and narrowing of the airways encourages impaction of particles. 
The larger the particle size, the greater the velocity of incoming air, while 
the greater the bend angle of bifurcations and the smaller the airway radius, the 
Aerosols as Drug Carriers 379 
greater the probability of deposition by impaction.66 Drug particles are known to 
be hygroscopic and grow in size in high humidity environments, such as the lung 
which has a relative humidity of approximately 99.5%. The addition and removal of 
water can significantly affect the particle size and thus deposition of a hygroscopic 
aerosol.67 A hygroscopic aerosol that is delivered at relatively low temperature and 
humidity into one of high humidity and temperature would be expected to increase 
in size when inhaled into the lung. The rate of growth is a function of the initial 
diameter of the particle, with the potential for the diameter of fine particles less 
than 1 /xm to increase 5-fold, compared with 2 to 3-fold for particles greater than 
2 /xm.68 The increase in particle size above the initial size should affect the amount 
of drug deposited, and particularly, the distribution of the aerosolized drug within 
the lung. Ferron and colleagues have predicted that for initial sizes between 0.7 /xm 
and 10 /xm, total deposition of hygroscopic aerosols increases by a factor of 2.69 
For particles with an initial size of 1 /xm, Xu and Yu were able to predict changes 
in the distribution pattern due to particle growth.70 The calculations showed a 
shift from deposition due to sedimentation to primarily impaction on more central 
airways.69 
5.2. Lung clearance mechanisms 
Once deposited in the lungs, inhaled drugs are either cleared from the lungs, 
absorbed into the circulatory or lymphatic systems, or metabolized. Drug particles 
deposited in the conducting airways are primarily removed through mucociliary 
clearance, and to a lesser extent, are absorbed through the airway epithelium into 
the blood or lymphatic system. Ciliated epithelium extends from the trachea to 
the terminal bronchioles. The airway epithelial goblet cells and submucosal glands 
secrete mucus forming a two-layer mucus blanket over the ciliated epithelium: a 
low-viscosity periciliary or sol layer covered by a high-viscosity gel layer. Insoluble 
particles are trapped in the gel layer and moved towards the pharynx (and 
ultimately to the gastrointestinal tract) by the upward movement of mucus generated 
by the metachronous beating of cilia. In the normal lung, the rate of mucus 
movement varies with the airway region and is determined by the number of 
ciliated cells and their beat frequency. Movement is faster in the trachea than in 
the small airways, and is affected by factors influencing ciliary functioning and 
the quantity and quality of the mucus.40'71 For normal mucociliary clearance to 
occur, airway epithelial cells must be intact, ciliary structure and activity normal, 
the depth and chemical composition of the sol layer optimal, and the rheology of 
the mucus within the physiological range. Mucociliary clearance is impaired in 
lung diseases such as immotile cilia syndrome, bronchiectasis, cystic fibrosis and 
asthma.72 In immotile cilia syndrome and bronchiectasis, the ciliary function can be 
380 Labiris, Bosco & Dolovich 
either impaired or nonexistent. In cystic fibrosis, the ciliary structure and function 
are normal, however, the copious amounts of thick, tenacious mucus present in 
the airways impairs their ability to clear the mucus effectively73 In these diseases, 
clearance of aerosolized drugs deposited in the conducting airways is generally 
decreased and secretions are cleared from the lung by cough.74-76 
In addition to mucociliary clearance, soluble particles can also be removed 
by absorptive mechanisms in the conducting airways.77 Lipophilic molecules pass 
easily through the airway epithelium via passive transport. Hydrophilic molecules 
cross via extracellular pathways such as tight junctions or by active transport via 
endocytosis and exocytosis.78 From the submucosal region, particles are absorbed 
either into systemic circulation, bronchial circulation or lymphatic systems. 
Drugs deposited in the alveolar region may be phagocytosed and cleared 
by alveolar macrophages or absorbed into the pulmonary circulation. Alveolar 
macrophages are the predominant phagocytic cell for the lung defense against 
inhaled microorganisms, particles and other toxic agents. There are approximately 
5 to 7 alveolar macrophages per alveolus in the lungs of healthy, non-smokers.79 
Macrophages phagocytose insoluble particles deposited in the alveolar region are 
either cleared by the lymphatic system or moved into the ciliated airways along currents 
in alveolar fluid and then cleared via the mucociliary escalator.65 This process 
can take weeks or months to complete.7 As discussed above, soluble drug particles 
deposited in the alveolar region can be absorbed into the systemic circulation. The 
pulmonary epithelium appears to be more resistant to soluble particle transport 
than to the endothelium or the interstitium.42 
The lung-blood barrier may behave as a molecular sieve, allowing the passage 
of small solutes but restricting the passage of macromolecules. Conhaim and colleagues 
proposed that the lung barrier was best fitted to a three pore size model, 
including a small number (2%) of large-sized pores (400 nm pore radius), 30% of 
medium-sized pores (40 nm radius) and 68% of small-sized pores (1.3 nm).80 
The rate of protein absorption from the alveoli is size dependent. Effros and 
Mason demonstrated an inverse relationship between alveolar permeability and 
molecular weight.42 In rats, after intratracheal instillation of DDAVP (1-desamino- 
8-D-arginine vasopressin) (raw = 1.1 kDa), peak serum DDAVP levels occurred 
at 1 hr compared with 16 to 24hrs after the intratracheal instillation of albumin 
(mw = 67 kDa).43 However, some proteins are cleared from the lung more rapidly 
than expected for their size. After intratracheal instillation or aerosolization of 
human growth hormone (mw = 22 kDa), peak serum levels were observed between 
0.5 to 4 hrs, indicating a rapid, saturable clearance from the lung that is suggestive 
of receptor-mediated endocytosis.65 Vasoactive intestinal polypeptide (VIP) 
is believed to be completely degraded during the passage across the pulmonary 
epithelium and into the bloodstream.81 
Aerosols as Drug Carriers 381 
Nanoparticles can pass rapidly into the systemic circulation. The distribution 
of radioactivity, after the inhalation of a 99mTechnetium (Tc)-labeled ultrafine carbon 
particles (5 to 10 nm), was detected in the blood one min post-inhalation and 
peaked between 10 and 20 min. This blood radioactivity level was sustained up to 
60 min. 8% of the initial lung radioactivity was measured in the liver 5 min postadministration 
and remained stable over time. The rapidity of the appearance of 
radioactivity systemically makes the translocation from the lung unlikely due to 
phagocytosis, by macrophages or endocytosis by epithelial and endothelial cells, 
but by passive diffusion.82 
6. Nanoparticle Formulations for Inhalation 
Delivery of nano-sized aerosols to the lung may result in very little drug being 
deposited in the lung. The majority of particles <500nm inhaled will not have 
enough residence time in the lung to deposit, and therefore will be exhaled (Fig. 1). 
However, if the nanoparticles were delivered in larger carrier particles, they could 
be sufficiently deposited in the lung. The carrier particle would dissolve after contact 
with the lung surface fluid, releasing the nanoparticle at the target tissue or cells. 
Sham and colleagues demonstrated that nanoparticles (173 to 242 nm) could 
be delivered into the lung in larger respirable lactose carrier particles produced 
by spray-drying.83 The dry powder containing the nanoparticles had a MMAD of 
3.0 /xm. pMDI formulations are typically micronized drugs in the 2 to 3 /xm range 
suspended in a hydrofluoroalkane (HFA) propellant. Solution pMDI such as QVAR 
produce smaller drug particles on propellant evaporation, resulting in better deposition 
and distribution than a micronized formulation.84 However, for insoluble 
drug particles in the propellant, the efficiency of pMDI is limited. A study by Dickinson 
et al. proposed the use of nanoparticles suspended in propellant as a method 
of increasing the delivery efficiency of insoluble drugs in pMDIs.85 They produced 
hydrophilic nanoparticles using a reverse phase microemulsion technique that captures 
nanoparticles by snap freezing, followed by freeze-drying. The nanoparticles 
of pure drug (salbutamol) and the drug in a non-polymer matrix (lecithin-based), 
with and without lactose, were dispersed in HFA-227 and in aerosol performance 
assessed by cascade impaction. The size of the salbutamol nanoparticles ranged 
from 34 to 216 nm. Dispersion of the nanoparticles in a HFA-227:hexane (95:5 v/v) 
blend resulted in a homogeneous fine suspension that showed no signs of sedimentation 
or creaming over several months. Rapid release of salbutamol from 
the nanoparticle was observed (approximately 4 min) as expected from the large 
surface area of the particles and the high water solubility of the drug. A high 
fine particle fraction (ex-device, % < 5.8 /xm) of 58.3% to 65.5% and a low MMAD 
382 Labiris, Bosco & Dolovich 
(1.2 to 1.5 /u.m) were observed with the nanoparticle formulations. This data suggests 
that a high fraction of the nanoparticles would be distributed in the alveolar 
region of the lung and represents the best aerosol that can be produced using 
a pMDI. 
Budesonide is a potent corticosteroid used as an inhaled anti inflammatory 
agent to treat asthma. It is available as a dry powder inhaler and as a suspension for 
inhalation with a nebulizer. A new formulation for nebulization has been developed 
that contains nanocrystals of budesonide that give the suspension solution-like 
qualities.86 The particles are 75 to 300 nm in diameter, compared with 4400 nm 
for the marketed budesonide suspension (Pulmicort Respules, AstraZeneca). In 
a randomized crossover study, 16 healthy volunteers were given the nanocrystal 
budesonide formulation (0.5 mg and 1.0 mg doses), Pulmicort respules and placebo 
via nebulization using a Pari LC jet nebulizer. Nebulization times were shorter 
for the nanocrystal formulation, compared with Pulmicort respules (~7.1 min vs. 
8.7 min). Similar AUCs were observed with the formulations, suggesting similar 
pulmonary absorption. However, a higher Cmax (1212pg/mL vs. 662pg/mL) and 
shorter Tmax (8.4 min vs. 14.4 min) for nanocrystal budesonide compared with the 
same dose of Pulmicort, suggests a more rapid drug delivery or absorption with 
the nanocrystal formulation. 
6.1. Diagnostic imaging 
Radiolabeled nanoparticles have been used for many years in pulmonary ventilation 
studies.87 Ultrafine 99mTc labeled carbon particles (Technegas) is a relatively 
new advance in ventilation scintigraphy.88 Technegas (Vita Medical Ltd., Sydney 
Australia) consists of nanoparticles of carbon with a diameter of approximately 
5 nm, that behaves more like a 0.2 /tm particle.89 Technegas is generated by the 
electrostatic heating of a graphite crucible to 2500°C in which a saline solution 
of 99mTc-pertechnetate had been placed and dried. The aerosol is dispersed in a 
lead-lined chamber in an atmosphere of 100% argon gas that is then inhaled by 
the patient. It is deposited in alveoli by inhalation and distributes similarly as the 
inert gas radioisotopes. Once they are inhaled, the particles adhere to the alveolar 
structures without appreciable movement for at least 40 min.88 
Pulmonary delivery of nanoparticles is also being investigated for lymphoscintigraphy 
to assess the spread of or the staging of lung cancer. Lung cancer 
usually exhibits metastasis proliferation, spreading through the lymphatic system 
and the blood circulation. Lymphatic drainage is responsible for the alveolar 
clearance of the deposited particulates and drugs up to a certain particle diameter 
(500 nm).90 Thus, radiolabeled nanoparticles could be used to visualize the lymph 
nodes to determine the presence of tumors. 
Aerosols as Drug Carriers 383 
The lymphatic uptake of solid lipid nanoparticles has also been studied as 
an imaging method to stage lung cancer. The lipid nanoparticles were radiolabeled 
with the lipophilic tracer, D,L-hexamethylpropylene amine oxime (HMPAO), 
tagged with 99 m-Tc. The lipid nanoparticles were prepared by the melted homogenization 
method and had a mean diameter of 200 nm.90 The radiolabeled nanoparticles 
were aerosolized using an ultrasonic nebulizer and delivered to rats until 
200,000 cpm was achieved over the lung. After inhalation, the total activity in the 
lung was observed, followed by a fast clearance rate (ti/2 = lOmin) that decreases 
activity in the lung to 25% of the total dose. Asignificant uptake (16.7%) was detected 
in the regional lymph nodes during the first 45 to 60 min, suggesting that aerosol 
delivery to the lungs of solid lipid nanoparticles could be used as an effective colloidal 
carrier for lymphoscintigraphy. 
Drainage into the lymph nodes following the lung instillation of nanoparticles 
of insoluble iodinated CT x-ray contrast agents was studied in beagle dogs.91 
Nanoparticles of the contrast agent were prepared by microfluidization. A particle 
size of 150 to 200 nm was achieved. The nanoparticles were suspended in 2 different 
surfactant solutions. 1.5 mL of the suspension was instilled using a fiber optic bronchoscope 
at specific sites in the small airways and alveoli. The nanoparticles were 
transported from the lung to the draining lymph nodes, 6 to 9 days post instillation 
as visible on the CT radiographs. No adverse clinical signs were observed in the 
dogs. However, microscopic lung lesions were observed at the instillation sites for 
both formulations and vehicle. The lesions consisted of inflammatory infiltrates, 
mainly macrophages, in intra-alveolar, interstitial and perivascular locations. A 
few small sites had fibrosis and granulomatous nodules with the destruction of the 
lung parenchyma. The presence of foamy macrophages was observed in the lymph 
nodes. The microscopic findings suggest that instillation of these nanoparticles of 
contrast agent may be harmful to the lung. The authors suggested that administering 
the nanoparticles as an aerosol, rather than by instillation, would prevent high 
concentrations in focal areas believed to be responsible for these lesions. 
6.2. Vaccine delivery 
Mucosal vaccine administration is an attractive method of inducing an immune 
response, since many pathogens invade the body through mucosal surfaces in the 
nose, lung and gut. As it is the first contact point, the mucosa has developed barriers 
to protect the body. The mucosa associated lymphoid tissue (MALT) is one of these 
barriers. It contributes 80% of the immunocytes and secretes more immunoglobulins 
than any other organs in the body.92 Antigens are delivered locally in the respiratory 
tract to nasal-associated and bronchus-associated lymphoid tissues (NALT 
and BALT, respectively) and a mucosal immunity is induced. Using nanoparticles, 
384 Labiris, Bosco & Dolovich 
systemic immunity may also be induced. Several studies have investigated the use 
of nanoparticles as carriers for the nasal delivery of vaccines. Using tetanus toxoid as 
a model antigen, Vila and colleagues have studied the use of chitosan nanoparticles 
as well as polyethyleneglycol and polylactic acid (PEG-PLA) nanoparticles as nasal 
vaccine carriers.93,94 They compared PEG-PLA nanoparticles with PLA alone.94 
Tetanus toxoid was entrapped in the hydrophobic PLA core and protected from 
interacting with enzymes such as lysozymes, by a hydrophilic PEG coating. Upon 
incubation with lysozymes in vitro, PLA particles aggregate and do not reach the 
epithelium, whereas PEG-PLA nanoparticles remain stable and size unmodified. 
The nanoparticles were produced using a double emulsion technique. PEG-PLA 
tetanus toxoid nanoparticles had a similar diameter to the PLA particles (196 nm 
vs. 188 nm), but had a lower loading efficiency of 33.4% compared with 48.1 % with 
PLA. The IgG antibody response induced by PEG-PLA was superior at weeks 2 to 
24, after intranasal instillation of 30 fig of tetanus toxoid (10 fil per nostril) on days 1, 
8 and 15 in male BALB/c mice. In a similar study, the same group compared radiolabeled 
PEG-PLA, PEG-PLA with gelatin stabilizer to radiolabeled PLA encapsulated 
tetanus toxoid. They reported that 1 hr after intranasal administration, PEG-PLA 
nanoparticles produced a radioactivity level 10-fold higher in the blood than PLA 
which remained constant for 24 hrs. The radioactivity detected in the lymph nodes, 
lungs, liver and spleen was 3 to 6 fold higher for PEG-PLA than PLA nanoparticles 
24 hrs post instillation. The results of this work suggest that the PEG-PLA 
nanoparticles are partially taken up by the M cells of the NALT, as well as being 
transported to the submucosa and drained into the lymphatic system and blood 
stream.95 Recent work by the same group has investigated the potential use of chitosan 
nanoparticles for nasal administration of vaccines.93 Chitosan is a hydrophilic 
natural polysaccharide that is biodegradable and has mucoadhesive properties. The 
nanoparticles are formed spontaneously by adding the counter anion sodium TPP 
into the chitosan solution, without the use of energy sources or organic solvents 
required for the production of PEG-PLA nanoparticles. Again, using tetanus toxoid 
as the model antigen, the investigators studied the effect of chitosan dose (200 fig 
and 70 fig) and molecular weight (23,38 or 70 kDa) on the efficacy of the nanoparticles. 
The nanoparticles produced were 300 to 350 nm and had a positive surface 
charge (+40 mV). The loading efficiency of tetanus toxoid was 50 to 60%, irrespective 
of the molecular weight of chitosan. In vitro, the formulations exhibited a rapid 
release over the initial 2 hrs followed by a slow release for 16 days, with the greater 
initial release at lower molecular weights of chitosan. 30 or 10 fig of antigen (associated 
with 200 and 70 fig of chitosan) was given intranasally to BALB/c mice on 
days 1, 8 and 15. The IgG levels induced by the nanoparticles were significantly 
higher than those elicited by free tetanus toxoid. The response lasted for the 24 
weeks studied with the IgG titres increasing over time. Anti-tetanus IgA titers were 
detected in the saliva, bronchoalveolar and intestinal lavage fluids 24 weeks post 
Aerosols as Drug Carriers 385 
administration. The results were independent of the administered dose and were 
significantly higher for the nanoparticle than the free tetanus toxoid. 
Jung and colleagues evaluated tetanus toxoid-loaded polymer nanoparticles as 
potential nasal vaccine carriers in mice.96 The nanoparticles were produced with 
various diameters (100 nm, 500 nm) using a novel polyester, sulfobutylated poly 
(vinyl alcohol)-graft-poly(lactide-co-glycolide), SB(43)-PVAL-g-PLGA. The surface 
charge was —43 to 59 mV. Mice were immunized with tetanus toxoid nanoparticles 
or free toxoid in solution at weeks 1, 2 and 3, either by oral, intranasal or 
intraperitoneal administration. Four weeks after the first intranasal immunization, 
IgG and IgA titers were significantly higher than baseline. Oral immunization with 
the nanoparticles produced a weak IgG antibody response. Only 10% of the oral 
dose was administered to the nose (2.89 vs. 28.9 /u,g), however, intranasal immunization 
appeared to be more effective in inducting an immune response. Particle size 
had an effect on the titer levels. Particles > 1 /i.m did not induce an immune response, 
but no difference was observed between the 500 nm and 100 nm nanoparticles which 
both induced significantly levels of IgG and IgA. 
These studies suggest that nasal delivery of vaccines using biodegradable 
nanoparticles are a promising method of inducing mucosal and systemic immunity. 
6.3. Anti Tuberculosis therapy 
Intracellular bacterial infections caused by pathogens such as Mycobacterium tuberculosis 
are difficult to eradicate because they are generally inaccessible to free antibiotics. 
By loading antibiotics into nanoparticles, it is expected that delivery to the 
infected cells would improve since nanoparticles have been shown to localize preferentially 
in organs with high phagocytic activity and in circulating macrophages 
as well.97 The encapsulation of antibiotics has several advantages: (1) It modifies 
their pharmacokinetic characteristics by prolonging the antibiotics half-life and 
increasing the area under the concentration time curve (AUC), while decreasing its 
apparent volume of distribution. (2) It improves the targeting of the drug to the 
phagocytic cells. (3) It reduces toxicity of the antibiotics, such as the hepatotoxicity 
of anti tuberculosis drugs and the nephrotoxicity of aminoglycosides. Antibiotics 
encapsulated in nanoparticles have been shown to be superior at treating intracellular 
infections when administered intravenously. However, the pulmonary delivery 
of these nanoparticles have only been investigated recently. 
Although effective therapy for tuberculosis is available, treatment failure and 
drug resistance is typically the result of patient's noncompliance. To improve compliance, 
investigators have been studying ways to reduce the dosing frequency of 
the drugs. Poly (lactide-co-glycolide) (PLG) nanoparticles as an aerosolized sustained 
release formulation for anti tuberculosis drugs, isoniazid, rifampicin and 
pyrazinamide, has been investigated since pulmonary tuberculosis is the most 
386 Labiris, Bosco & Dolovich 
common form of the infection.98 The majority of the nanoparticles were 186 to 
290 nm in diameter. Drug encapsulation efficiency was 56.9% to 68%. Aerosolized 
nanoparticles had a MMAD of 1.88 //.m, with 96% of the particles in the respirable 
range (<6 /u,m). A single nebulization to guinea pigs resulted in sustained plasma 
drug concentrations for 6 to 8 days and in the lung for 11 days. The half-life and 
mean residence time of the drugs was significantly prolonged, compared with the 
oral free drugs. Nebulizing the nanoparticles every 10 days to guinea pigs infected 
with Mycobacterium tuberculosis resulted in no detectable bacilli in the lung after 5 
doses of treatment, compared with 46 daily doses of orally administered drug to 
achieve the equivalent efficacy. 
The use of lectin-based PLG nanoparticles as an aerosolized sustained release 
formulation of isoniazid, rifampicin and pyrazinamide has also been studied in 
guinea pigs." Mucoadhesive drug delivery systems such as chitosan have been 
previously investigated as a method of prolonging residence at a site of absorption. 
The main drawback of mucoadhesive systems is that its residence time is limited 
by the turnover time of the mucous gel layer, which is only a few hrs. Attaching the 
polymeric nanoparticles to cytoadhesive ligands such as lectins could prolong the 
duration of adhesion, thereby prolonging residence time. Lectins bind to epithelial 
surfaces via specific receptors. Wheat germ agglutinin (WGA) is the least immunogenic 
lectin and has known receptors on the alveolar epithelium as well as the 
intestinal wall. WGA lectin-PLG nanoparticles were prepared by a two-step carbodiimide 
procedure. Their size ranged from 350 to 400 nm with drug encapsulation 
efficiency between 54% and 66%. The nanoparticles were delivered via nebulization 
to guinea pigs. 88% of the aerosol was in the respirable range (<6/U,m) with 
a MMAD of 2.8/zm (GSD of 2.1). Three doses of nanoparticles were administered 
every 15 days for 45 days. The WGA-PLG nanoparticles resulted in a prolonged 
Tmax/ increased AUC and mean residence time after inhaled delivery. All three drugs 
were present in the lungs, liver and spleen at concentrations above the minimum 
inhibitory concentration 15 days post dosing, compared with orally-administered 
free drug. Chemotherapeutic studies in guinea pigs infected with Mycobacterium 
tuberculosis showed that 3 doses administered every 15 days for 45 days yielded 
undetectable mycobacterial colony forming units, which was only achievable with 
45 doses of the oral free drugs. The study results suggest that WGA-based PLG 
nanoparticles could be potential drug carriers for anti tuberculosis through aerosol 
delivery, reducing the drug dosing frequency. 
6.4. Gene therapy 
Pulmonary gene delivery and DNA vaccinations are attractive therapies for a variety 
of lung diseases such as cystic fibrosis, asthma, chronic obstructive pulmonary 
Aerosols as Drug Carriers 387 
disease, lung cancer and infections caused by Mycobacterium tuberculosis, influenza 
or SARS-associated coronavirus. Gene delivery requires carriers to transfer DNA 
into the nuclei of cells. There are two approaches for delivery: viral and non viral 
carriers. Viral delivery systems, although very efficient at transfection, are problematic 
due to their inherent immunogenicity. Non viral are safer but their transfection 
efficiency is low. Recently, biodegradable polymer-based nanoparticles have been 
investigated as a non viral pulmonary gene delivery system, taking advantage of 
their prolonged residence time in the lung and ability to be taken up by macrophages 
and dendritic cells, and to escape degradation by lysosomes. 
Asthma is characterized by elevated eosinophilic inflammation in the airway 
and increased airway hyperresponsiveness. Chronic inflammation can lead to structural 
damage and airway remodeling. IFN-y is a cytokine that promotes T-helper 
type 1 (Thl) responses which down regulates the Th2 immune responses present 
in asthma. Recombinant IFN-y has been shown to reverse inflammation in murine 
models of asthma. However, its short half-life and severe adverse effects at high 
doses have prevented its therapeutic use.100 An intra-nasal IFN-y gene therapy 
had been developed as an attempt to circumvent the drawbacks to its use. Kumar 
and colleagues studied the effects of a chitosan-IFN-y plasmid DNA nanoparticle 
in a BALB/c mouse model of allergic asthma (using ovalbumin-sensitization).101 
Mice treated with the chitosan nanoparticles exhibited a significantly lower airway 
hyperresponsiveness (to methacholine challenge), reduced number of eosinophils 
and a significant decrease in epithelial denudation, mucus cell hyperplasia and 
cellular infiltration. Production of IFN-y was increased post-treatment while IL-5 
and IL-4 and ovalbumin-specific IgE were reduced. Chitosan IFN-y nanoparticles 
induced IFN-y gene expression predominately in epithelial cells and worked within 
3 to 6 hrs after intranasal administration. 
Poly (D,L-lactide-co-glycolide (PLGA)-polyethyleneimine (PEI) nanoparticles 
are also being investigated for pulmonary gene delivery. PLGA had been extensively 
evaluated for its sustained-release profile and ability to be taken up by 
macrophages. PEI is a cationic polymer. Its high positive charge density suggests 
that it would be a promising candidate as a non viral vector.102 PLGA nanoparticles 
with PEI on their surface had a mean particle diameter between 207 and 231 nm, 
surface charge > 30 mV and a DNA loading efficiency of >99%. Internalization of 
the nanoparticles in the human airway submucosal epithelial cell line, Calu-3, was 
observed and DNA detected 6 hrs after administration. However, in vivo efficiency 
of this system still needs to be studied. 
Respiratory syncytial virus (RSV) infection is a major cause of respiratory tract 
infections and is associated with approximately 17000 deaths annually on a worldwide 
basis, with no anti viral therapy or vaccine available.103 RSV NS1 protein 
appears to antagonize the host type 1 interferon-mediated response. Zhang and 
388 Labiris, Bosco & Dolovich 
colleagues hypothesized that blocking the NS gene expression might inhibit RSV 
replication and thus provide effective antiviral therapy.104 Small interfering RNA 
(siRNA) targeting the NS1 gene (siNSl) were encapsulated in chitosan nanoparticles. 
BALB/c mice were intranasally treated with siNSl chitosan-nanoparticles 
before or after RSV infection. A significant decrease in virus titers in the lung was 
observed, in addition to a decrease in inflammation and airway hyperresponsiveness, 
compared with controls. The effect of siNSl lasted at least 4 days. The data 
show that siNSl nanoparticles may be a promising anti viral therapy against RSV 
infection. 
7. Conclusion 
Innovations in the biotechnology and pharmaceutical industries have led to novel 
approaches for delivering drugs more efficiently and to specific targets in the lung 
and the body. One of the growth areas is the development of nanoparticles as carriers 
of active pharmaceutical agents for diagnosis and treatment. 
Aerosol delivery systems, discussed at the beginning of this chapter, are the 
current technologies for delivering therapies to treat respiratory diseases and some 
systemic diseases. The accepted philosophy, and one based on sound in vitro and 
in vivo clinical data, is that the optimal size of aerosol needed to target the distal 
lung is of the order of 3 /u,m. This size is 10-100 times greater than the nanoparticles 
being considered in the design of agents including antibiotics, vaccines and gene 
therapies for inhaled delivery. Novel techniques and formulations are being studied 
to produce successful vehicles for delivering these types of products in vivo. Positive 
outcomes using animal models to test these new aerosol formulations have been 
reported. However, clinical studies still need to be conducted to determine their 
efficacy in humans. 
As with any new technology, there will be benefits and risks associated with 
its use. The use of nanotechnology to provide improved targeting of drugs via 
the inhaled route is an exciting development that has the potential to yield novel 
treatments for many diseases in the near future. 
References 
1. Horsfield K, Dart G, Olson DE, Filley GF and Cumming G (1971) Models of the human 
bronchial tree. / Appl Physiol 31:207-217. 
2. Kreuter J (1994) Drug targeting with nanoparticles. Eur J Drug Met Pharmacol 
19:253-256. 
3. Ghirardelli R, Bonasoro F, Porta C and Cremaschi D (1999) Identification of particular 
epithelial areas and cells that transport polypeptide-coated nanoparticles in the nasal 
respiratory mucosa of the rabbit. Biochim Biophys Acta 12:1-2. 
Aerosols as Drug Carriers 389 
4. Huang M, Ma Z, Khor E and Lim LY (2002) Uptake of FITC-chitosan nanoparticles in 
A549 cells. Pharm Res 19:1488-1494. 
5. Russell-Jones GJ, Veitch H and Arthur L (1999) Lectin-mediated transport of nanoparticles 
across Caco-2 and OK cells. Int f Pharm 190:165-174. 
6. Brigger I, Dubernet C and Couvreur P (2002) Nanoparticles in cancer therapy and 
diagnosis. Adv Drug Del Rev 54:631-651. 
7. Martonen TB (1993) Mathematical model for the selective deposition of inhaled pharmaceuticals. 
/ Pharm Sci 82:1191-1199. 
8. O'Callaghan C and Barry PW (1997) The science of nebulized drug delivery. Thorax 52 
(Suppl 2):s31-s44. 
9. Denyer J, Dyche T, Nikander K, Newman SP, Richards J and Dean A (1997) Halolite a 
novel liquid drug aerosol delivery system. Thorax 52(suppl 6):208. 
10. Dolovich M (1999) New propellant-free technologies under investigation. / Aerosol Med 
12(Suppl I):s9-sl7. 
11. Smaldone GC, Agosti J, Castillo R, Cipolla D and Blanchard JD (1999) Deposition of 
radiolabelled protein from AERx in patients with asthma. / Aerosol Med 12:98. 
12. Newman SP and Clarke SW (1992) Inhalation devices and techniques, in Clark TJH, 
Godfrey S and Lee TH (eds.), Asthma. Chapman & Hall, London, pp. 469-^55. 
13. Newman SP, Pavia D, Moren F, Sheahan NF and Clarke SW (1981) Deposition of pressurized 
aerosols in the human respiratory tract. Thorax 36:52-55. 
14. Crompton GK (1982) Problems patients have using pressurized aerosol inhalers. 
European Journal of Respiratory Diseases — Supplement 119:101-104. 
15. Bennett WD and Smaldone GC (1987) Human ventilation in the peripheral air-space 
deposition of inhaled particles. Am } Physiol 62:1603-1610. 
16. Dolovich M, Ryan G and Newhouse MT (1981) Aerosol penetration into the lung. 
Influence on airway responses. Chest 80 (suppl)(6):834-836. 
17. Newman SP, Pavia D, Garland N and Clarke SW (1982) Effects of various inhalation 
modes on the deposition of radioactive pressurized aerosols. Eur } Respir Dis 
63(Suppl. 119):57-65. 
18. Pavia D, Thomson ML, Clarke SW and Shannon HS (1977) Effects of lung function and 
mode of inhalation on penetration of aerosol into the human lung. Thorax 32:194-197. 
19. Dolovich M, Ruffin RE, Roberts R and Newhouse MT (1981) Optimal delivery of 
aerosols from metered dose inhalers. Chest 80:911-915. 
20. Dolovich MB (1995) Characterization of medical aerosols: Physical and clinical requirements 
for new inhalers. Aerosol Sci Technol 22:392-399. 
21. Dolovich M, Chambers C, Mazza M and Newhouse MT (1992) Relative efficiency of 
four metered dose inhaler (MDI) holding chambers (HC) compared to albuterol MDI. 
J Aerosol Med 5:307. 
22. Dolovich M, Ruffin R, Corr D and Newhouse MT (1983) Clinical evaluation of the 
Aerochamber: A simple demand inhalation MDI delivery device. Chest 84:36-41. 
23. Newman SP, Moren F, Trofast E, Talaee N and Clarke SW (1989) Deposition and clinical 
efficacy of terbutaline sulphate from Turbohaler, a new multi-dose inhaler. Eur Respir } 
2:247-252. 
390 Labiris, Bosco & Dolovich 
24. Pedersen S (1996) Inhalers and nebulizers: Which to choose and why. Respiratory 
Medicine 90:69-77. 
25. Newhouse MT and Kennedy A (2000) Condensation due to rapid, large temperature 
(t) changes impairs aerosol dispersion from Turbuhaler(T). Am } Respir Cell Mol Biol 
161(3):A35. 
26. Newhouse MT and Kennedy A (2000) Inspiryl Turbuhaler (ITH) DPI vs. Ventolin 
MDI + Aerochamber (AC): Aerosol dispersion at high and low flow and relative humidity/
temperature (RH/T) in vitro. Am } Respir Crit Care Med 161(3):A35. 
27. Borgstrom L, Bondesson E, Moren F, Trofast E and Newman SP (1994) Lung deposition 
of budesonide inhaled via Turbuhaler: A comparison with terbutaline sulphate in 
normal subjects. Euro Respir } 7:(1) 69-73. 
28. Richards R and Saunders M (1993) Need for a comparative performance standard for 
dry powder inhalers. Thorax 48:1186-1187. 
29. Everard ML, Devadason SG and Le Souef PN (1997) Flow early in the inspiratory 
manoeuvre affects the aerosol particle size distribution from a Turbuhaler. Respir Med 
91:624-628. 
30. Newman SP, Hollingworth AandClarkAR(1994)Effectof different modes of inhalation 
on drug delivery from a dry powder inhaler. Int} Pharma 102:127-132. 
31. Newman SP, Moren F, Trofast E, Talaee N and Clarke SW (1991) Terbutaline sulphate 
Turbuhaler: Effect of inhaled flow rate on drug deposition and efficacy. Int J Pharma 
74:209-213. 
32. Clark AR and Hollingworth AM (1993) The relationship between powder inhaler resistance 
and peak inspiratory conditions in healthy volunteers — Implications for in vitro 
testing.} Aerosol Med 6(2):99-110. 
33. Svartengren K, Lindestad PA, Svartengren M, Philipson K, Bilin G and Camner P 
(1995) Added external resistance reduces oropharyngeal deposition and increases 
lung deposition of aerosol particles in asthmatics. Am J Respir Crit Care Med 
152:32-37. 
34. Melchor R, Biddiscombe MF, Mak VH, Short MD and Spiro SG (1993) Lung deposition 
patterns of directly labelled salbutamol in normal subjects and in patients with 
reversible airflow obstruction. Thorax 48(5)506-511. 
35. Vid gren M, Paronen P, Vidgren P, Vainio P and Nuutinen J (1990) In vivo evaluation of 
the new multiple dose powder inhaler and the Rotahaler using the gamma scintigraphy. 
Acta Pharma Nordica 2(1):3-10. 
36. Borgstrom L and Newman SP (1993) Total and regional lung deposition of terbutaline 
sulphate inhaled via a pressurised MDI or via Turbuhaler. Int ] Pharma 97:47-53. 
37. Ruffin RE, Dolovich MB, Wolff RK and Newhouse MT (1978) The effects of preferential 
deposition of histamine in the human airway. Am Rev Respir Dis 117(3) :485-492. 
38. Lourenco RV and Cotromanes E (1982) Clinical aerosols. I. Characterization of aerosols 
and their diagnositic uses. Arch Intern Med 142:2163-2172. 
39. Heyder J. (1982) Particle transport onto human airway surfaces. Eur } Respir Dis 
63(Suppl 119):29-50. 
Aerosols as Drug Carriers 391 
40. Brain JD and Blanchard JD (1993) Mechanisms aerosol deposition and clearance, in 
Moren F, Newhouse MT and Dolovich MB (eds.), Aerosols in Medicine. Principles, Diagnosis 
and Therapy. Elsevier Science Publishers, New York, pp. 117-156. 
41. Gerrity TR (1990) Pathophysiological and disease constraints on aerosol deposition, in 
Byron PR (ed.), Respiratory Drug Delivery. Boca Raton, CRC Press, Inc., Florida, 1990: 
1-38. 
42. Effros RM and Mason GR (1983) Measurements of pulmonary epithelial permeability 
in vivo. Am Rev Respir Dis 127(suppl):s59-s66. 
43. Folkesson HG, Westrom BR and Karlsson BW (1990) Permeability of the respiratory 
tract to different-sized macromolecules after intratracheal instillation in young and 
adult rats. Acta Physiol Scand 139:347-354. 
44. Patton JS (1996) Mechanisms of macromolecule absorption by the lungs. Adv Drug Del 
Rev 19:3-36. 
45. Carstairs JR, Nimmo AJ and Barnes PJ (1985) Autoradiographic visualization of betaadrenoceptor 
subtypes in human lung. Am Rev Respir Dis 132:541-547. 
46. Mak JCW and Barnes PJ (1990) Autoradiographic visualization of muscarinic receptor 
subtypes in human and guinea pig lung. Am Rev Respir Dis 141:1559-1568. 
47. Rees PJ, Clark TJ and Moren F (1982) The importance of particle size in response 
to inhaled bronchodilators. European Journal of Respiratory Diseases — Supplement 
119:73-78. 
48. Zanen P, Go LT and Lammers J-WJ (1994) The optimal particle size for /^-adrenergic 
aerosols in mild asthmatics. Int J Pharma 107:211-217. 
49. Zanen P, Go LT and Lammers J-WJ (1996) Optimal particle size for Pi agonist 
and anticholinergic aerosols in patients with severe airflow obstruction. Thorax 
51:977-980. 
50. Baskin ML, Abd AG and Ilowite JS (1990) Regional deposition of aerosolized pentamidine. 
Effects of body position and breathing pattern. Ann Int Med 113:677-683. 
51. Gerrity TR, Garrard CS and Yeates DB (1981) Theoretical analysis of sites of aerosol 
deposition in the human lung. Chest 80(suppl 6):898-901. 
52. Baltimore RS, Christie CDC and Walker Smith GJ (1989) Immunohistopathologic localization 
of Pseudomonas aeruginosa in lungs from patients with cystic fibrosis. Implications 
for the pathogenesis of progressive lung deterioration. Am Rev Respir Dis 
140:1650-1661. 
53. Potts SB, Roggli VL and Spock A (1995) Immunohistologic quantification of 
Pseudomonas aeruginosa in the tracheobronchial tree from patients with cystic fibrosis. 
Pediatric Pathol Lab Med 15:707-721. 
54. Alderson PO, Seeker-Walker RH, Stominger DB et al. (1974) Pulmonary deposition of 
aerosols in children with cystic fibrosis. / Pediatr 84:479^84. 
55. Ilowite JS, Gorvoy JD and Smaldone GC (1987) Quantitative deposition of aerosolized 
gentamicin in cystic fibrosis. Am Rev Respir Dis 136:1445-1449. 
56. Anderson PJ, Blanchard JD, Brain JD, Feldman HA, McNamara JJ and Heyder J (1989) 
Effect of cystic fibrosis on inhaled aerosol boluses. Am Rev Respir Dis 140:1317-1324. 
392 Labiris, Bosco & Dolovich 
57. Wolff RK (1998) Safety of inhaled proteins for therapeutic use. / Aerosol Med 
11(4):197-219. 
58. Farr ST, Gonda I and Licko V (1998) Physicochemical and physiological factors influencing 
the effectiveness of inhaled insulin, in Dalby RN, Byron PR and Farr ST (eds.), 
in Respiratory Drug Delivery VI, Interpharm Press Inc., Buffalo Grove, IL, pp. 25-33. 
59. Laube BL, Georgopoulos A and Adams GKI (1993) Preliminary study of the efficacy 
of insulin aerosol delivered by oral inhalation in diabetic patients. JAMA 269(16): 
2106-2109. 
60. Jendle JH and Karlberg BE (1996) Effects of intrapulmonary insulin in patients with 
non-insulin-dependent diabetes. Scand } Clin Lab Invest 56:555-561. 
61. Jendle JH and Karlberg BE (1996) Intrapulmonary administration of insulin to healthy 
volunteers. / Int Med 240:93-98. 
62. Laube BL, Benedict GW and Dobs AS (1998) Time to peak insulin level, relative bioavailability, 
and effect of size of deposition of nebulized insulin in patients with noninsulindependent 
diabetes mellitus. / Aerosol Med 11 (3): 153-173. 
63. Byron PR and Patton JS (1994) Drug delivery via the respiratory tract. / Aerosol Med 
7(l):49-75. 
64. Ma JKH, Bhat M and Rojanasakul Y (1996) Drug metabolism and enzyme kinetics in 
the lung, in: Lenfant C (ed.), Inhalation Aerosols. Physical and Biological Basis for Therapy. 
Marcel Dekker: New York, NY, 94:155-195. 
65. Folkesson HG, Matthey MA, Westrom BR, Kim KJ, Karlsson BW and Hastings RH 
(1996) Alveolar epithelial clearance of protein. / Appl Physiol 80(5): 1431-1445. 
66. Newman SP (1985) Aerosol deposition considerations in inhalation therapy. Chest 
88(suppl):152s-160s. 
67. Phipps PR, Gonda I, Anderson SD, Bailey D and Bautovich G (1994) Regional deposition 
of saline aerosols of different tonicities in normal and asthmatic subjects. Euro Respir J 
7(8):1474-1482. 
68. Swift DL (1980) Aerosols and humidity therapy: Generation and respiratory deposition 
of therapeutic aerosols. Am Rev Respir Dis 122:71-91. 
69. Ferron GA, Oberdorster G and Henneberg R (1989) Estimation of the deposition of 
aerosolized drugs in the human respiratory tract due to hygroscopic growth. / Aerosol 
Med 2(3):271-284. 
70. Xu GB and Yu CP. (1985) Theoretical lung deposition of hygroscopic NaCl aerosols. 
Aerosol Sci Technol 4:455-461. 
71. Smaldone GC, Perry RJ, Bennett WD, Messina MS, Zwang J and Ilowite J (1988) Interpretation 
of "24 hour lung retention" in Studies of Mucociliary Clearance. / Aerosol Med 
1:11-20. 
72. Houtmeyers E, Gosselink R, Gayan-Ramirez G and Decramer M (1999) Regulation of 
mucociliary clearance in health and disease. Eur Respir } 13:1177-1188. 
73. Rossman CM, Lee RMKW, Forrest JB and Newhouse MT (1984) Nasal ciliary ultrastructure 
and function in patients with primary ciliary dyskinesia compared with that 
in normal subjects and in subjects with various respiratory diseases. Am Rev Respir Dis 
129(1):161-167. 
Aerosols as Drug Carriers 393 
74. Rossman CM, Waldes OR, Sampson D and Newhouse MT (1982) Effect of chest physiotherapy 
on the removal of mucus in patients with cystic fibrosis. Am Rev Respir Dis 
126:131-135. 
75. Robinson M (2002) Bye PTB. Mucociliary clearance in cystic fibrosis. Pediatric Pulmonol 
33:293-306. 
76. Isawa T, Teshima T, Hirano T, Anazawa Y, Miki M, Konno K and Motomiya M (1990) 
Mucociliary clearance and transport in bronchiectasis: Global and regional assessment. 
JNucl Med 31:543-548. 
77. Edsbacker S (2002) Uptake, retention and biotransformation of corticosteroids in the 
lung and airways, in Schleimer RP, O'Byrne PM, Szefler SJ and Brattsand R (eds.), 
Inhaled Steroids in Asthma: Optimizing Effects in the Airways, Marcel Dekker, Inc., 
New York, pp. 213-246. 
78. Summers QA. (1991) Inhaled drugs and the lung. Clin Exp Allergy 21:259- 
268. 
79. Stone KC, Mercer RR, Gehr P, Stockstill B and Crapo JD (1992) Allometric relationship 
of cell numbers and size in the mammalian lung. Am } Respir Cell Mol Biol 
6:235-243. 
80. Conhaim RL, Eaton A, Staub NC and Heath TD (1988) Equivalent pore estimate 
for the alveolar-airway barrier in isolated dog lung. / Appl Physiol 64:1134- 
1142. 
81. Barrowcliffe MPA, Jones JG and Sever PS (1986) Pulmonary clearance of vasoactive 
intestinal peptide. Thorax 41:88-93. 
82. Nemmar A, Hoet PHM, Vanquickenborne B, Dinsdale D, Thomeer M, Hoylaerts MF, 
Vanbilloen H, Mortelmans L and Nemery B (2002) Passage of inhaled particles into the 
blood circulation in humans. Circulation 105:411-414. 
83. Sham JO-H, Zhang Y, Finlay WH, Roa WH and Lobenberg R (2004) Formulation and 
characterization of spray-dried powders containing nanoparticles for aerosol delivery 
to the lung. Int J Pharma 269:457-467. 
84. Leach CL, Davidson PJ and Boudreau RJ (1998) Improved airway targeting with 
the CFC-free HFA-beclomethasone metered dose inhaler compared with CFCbeclomethasone. 
Eur Respir } 12:1346-1353. 
85. Dickinson PA, Howells SW and Kellaway IW (2001) Novel nanoparticles for pulmonary 
drug administration. / Drug Targ 9(4):295-302. 
86. Kraft WK, Steiger B, Beussink D, Quiring JN, Fitzgerald N, Greenberg HE and Waldman 
SA (2004) The pharmacokinetics of nebulized nanocrystal budesonide suspension in 
healthy volunteers. / Clin Pharma 44(l):67-72. 
87. Kramer ELandDivgiCR(1991)Pulmonary applications of nuclear medicine. Clin Chest 
Med 12(l):55-75. 
88. Satoh K, Takahashi K, Sasaki M, Kobayashi T, Honjo N, Ohkawa M, Tanabe M, Fujita J 
and Miyawaki H (1997) Comparison of 99mTc-Technegas SPECT with 133Xe dynamic 
SPECT in pulmonary emphysema. Ann Nuclear Med 11:201-206. 
89. Isawa T, Teshima T, Anazawa Y, Miki M and Motomiya M (1991) Technegas for inhalation 
lung imaging. Nucl Med Commun 12:47-55. 
394 Labiris, Bosco & Dolovich 
90. Videira MA, Botelho MF, Santos AC, Gouveia LF, Pedrosos De Lima JJ and Almeida AJ 
(2002) Lymphatic uptake of pulmonary delivered radiolabelled solid lipid nanoparticles. 
/ Drug Targ 10(8):607-613. 
91. Mclntire GL, Bacon ER, Toner JL, Cornacoff JB, Losco PE, Illig KJ, Nikula KJ, 
Muggenburg BA and Ketai L (1998) Pulmonary delivery of nanoparticles of insoluble, 
iodinated CT X-ray contrast agents to lung draining lymph nodes in dogs. J Pharma Sci 
87(11):1466-1470. 
92. Lamm ME (1997) Interactions of antigens and antibodies at mucosal surfaces. Ann Rev 
Microbiol 51:311-340. 
93. Vila A, Sanchez A, Janes K, Behrens I, Kissel T, Vila Jato J and Alonso MJ (2004) Low 
molecular weight chitosan nanoparticles as new carriers for nasal vaccine delivery in 
mice. Eur J Pharmaceut Biopharmeut 57:123-131. 
94. Vila A, Sanchez A, Evora C, Soriano I, Vila Jato J and Alonso MJ (2004) PEG-PLA 
nanoparticles as carriers for nasal vaccine delivery. / Aerosol Med 17(2):174-185. 
95. Tobio M, Gref R, Sanchez A, Langer R and Alonso MJ (1998) Stealth PLAPEG 
nanoparticles as protein carriers for nasal administration. Pharma Res 15(2): 
270-275. 
96. Jung T, Kamm W, Breitenbach A, Hungerer K-D, Hundt E and Kissel T (2001) 
Tetanus toxoid loaded nanoparticles from sulfobutylated poly(vinyl alcohol)-graftpoly(
lactide-co-glycolide): Evaluation of antibody response after oral and nasal application 
in mice. Pharma Res 18(3):352-360. 
97. Pinto-Alphandary H, Andremont A and Couvreur P (2000) Trgeted delivery of antibiotics 
using liposomes and nanoparticles: Research and applications. Int J Antimicrobial 
Agents 13(3):155-168. 
98. Pandey R, Sharma A, Zahoor A, Sharma S, Khuller GK and Prasad B (2003) Poly (DLlactide-
co-glycolide) nanoparticle-based inhalable sustained drug delivery system for 
experimental tuberculosis. / Antimicrob Chemother 52(6):981-986. 
99. Sharma A, Sharma S and Khuller GK (2004) Lectin-functionalized poly (lactide-coglycolide) 
nanoparticles as oral/aerosolized antitubercular drug carriers for treatment 
of tuberculosis. / Antimicrob Chemother 54(4):761-766. 
100. Yoshida M, Leigh R, Matsumoto K, Wattie J, Ellis R and O'Byrne PM (2002) Effect of 
interferon-gamma on allergic airway responses in interferon-gamma-deficient mice. 
Am } Respir Crit Care Med 166:451^56. 
101. Kumar M, Kong X, Behera AK, Hellermann GR, Lockey RF and Mohapatra SS (2003) 
Chitosan IFN-y-pDNA nanoparticle (CIN) therapy for allergic asthma. Gen Vacc Ther 
1:3-11. 
102. Bivas-Benita M, Romeijn S, Junginger HE and Borchard G (2004) PLGA-PEI nanoparticles 
for gene delivery to pulmonary epithelium. Eur J Pharmaceut Biopharmeut 58:1-6. 
103. Thompson WW, Shay DK, Weintraub E, Brammer L, Cox N, Anderson LJ and Fukuda K 
(2003) Mortality associated with influenza and respiratory syncytial virus in the United 
States. JAMA 289(2):179-186. 
104. Zhang W, Yang H, Kong X, Mohapatra S, Juan-Vergara HS, Hellermann G, Behera S, 
Singam R, Lockey RF and Mohapatra SS (2005) Inhibition of respiratory syncytial virus 
Aerosols as Drug Carriers 395 
infection with intranasal siRNA nanoparticles targeting the viral NSl gene. Nat Med 
ll(l):56-62. 
105. Dolovich MB and Newhouse MT. (1993) Aerosols. Generation, methods of administration, 
and therapeutic applications in asthma, in Middleton E Jr, Reed CE, Ellis EF, 
Adkinson NF Jr, Yunginger JW and Busse WW (eds.), Allergy. Principles and Practice, 
Mosby — Year Book, Inc., St. Louis, 712-739. 
This page is intentionally left blank
18 
Magnetic Nanoparticles as 
Drug Carriers 
Urs O. Hafeli and Mathieu Chastellain 
Magnetic nanoparticles possess many characteristics that make them promising 
as drug carriers and for use in biomedical applications. They can be attracted 
or magnetically guided by strong magnetic fields, thus acting as drug carriers. 
They can also be used for hyperthermia applications, due to the heat they produce 
in an alternating magnetic field. The resulting temperature increase can be 
used to modify or inhibit specific cell activities locally, or even to release drugs in a 
precisely controlled, temperature-increase activated manner. Magnetic nanoparticles 
can also serve as contrast agents for diagnostic applications such as magnetic 
resonance imaging. 
1. Introduction 
Magnetic nanoparticles occur frequently in nature. They are found not only in the 
mineral world but also in living organisms. Well known examples are magnetotactic 
bacteria, which are believed to navigate the waters they live in, by using internal 
magnetic crystals aligned in chains that function as a compass. Higher forms of life, 
such as humans, also employ iron as an essential metal. In order to ensure a constant 
supply of iron, the body stores it within the well-defined protein shell ferritin as a 
5 to 7 nm hydrous ferric oxide nanoparticle.1,2 
The use of magnetic powders in medical applications was already conceptualized 
by ancient Greek and Roman scientists.3 However, magnetic nanoparticles 
have only been used since the mid 1970s in the area of biological and medical 
sciences.4 A wide range of in vivo as well as in vitro applications have been or are 
397 
398 Hafeli & Chastellain 
currently being developed.5~u These applications include magnetic drug delivery, 
magnetic fluid hyperthermia, magnetic cell separation and extraction when an 
external magnetic field is applied, and contrast enhancement for diagnostic imaging 
procedures, as the magnetic nanoparticles' own magnetic field influences their 
surrounding. From a practical point of view, magnetic nanoparticles are thus versatile 
tools that enhance yields for many in vitro processes such as cell purifications. 
In addition, in general, no invasive procedures are required when they are used for 
in vivo therapies. 
2. Definitions 
The development of nanoparticles for biomedical applications requires contributions 
from the basic to the medical sciences. Such interdisciplinary interactions 
can sometimes lead to communication problems. For example, the term magnetic 
nanoparticle has a different meaning for a physician and a biochemist, and might 
have no meaning at all for a physicist. For this reason, definitions satisfying all 
partners involved in the present research field are required. With the following 
simplified definitions, we attempt to provide a universal starting point. 
2.1. Properties of magnetic ma terials 
The magnetic properties of materials are mainly related to electrons, with all materials 
showing some kind of magnetic behavior. Materials can be classified according 
to their response to external magnetic solicitations. Magnetic susceptibility is 
defined by the initial slope of the magnetic curve, presenting the magnetization 
"M" (response) as a function of an applied magnetic field "H" (solicitation) (Fig. 1). 
The observed behavior of different materials can be explained in terms of their 
magnetic structure at the atomic level, and can be summarized as diamagnetism, 
paramagnetism, and ferromagnetism. Diamagnetic materials consist of atoms with 
no net magnetic moment. Nevertheless, they tend to oppose any external magnetic 
field change due to induced dipoles in the material. For this reason, they are 
characterized by a slight negative magnetic susceptibility. Paramagnetic materials 
are made of atoms showing a net magnetic moment. The random orientation 
of these moments is responsible for a slight positive magnetic susceptibility and 
no magnetization remains when the external magnetic field is switched off (see 
Fig. 1). Ferromagnetic materials react strongly to external magnetic fields, unlike 
dia- and para-magnetic materials. They can be viewed as paramagnetic materials 
with an organized domain structure (see Fig. 1). Within a domain, all atomic 
magnetic moments are parallel. When submitted to an external magnetic field, the 
different domains, initially in a random orientation, tend to align according to the 
Magnetic Nanoparticles as Drug Carriers 399 
Temperature 
V increase ... ^'" Size 
increase 
paramagnetism ferro- or ferri-magnetism superparamagnetism *m VX-*; 
M i 
- • H 
Thermal activation 
^ atomic magnetic moment 
M magnetisation 
H applied magnetic field 
• H 
MR remanent magnetisation 
Hc coercive field 
Ms saturation magnetisation 
Fig. 1. Atomic magnetic moment structure (upper drawings) and corresponding magnetization 
curves (lower graphs). Paramagnetic materials show random atomic moment 
orientation which is responsible for their weak response to magnetic solicitations and no 
remanence. A typical ferro- or ferri-magnetic material shows a characteristic domain structure 
with associated hysteresis magnetization curve. Superparamagnetic materials present 
a thermally induced oscillating magnetic moment and a strong magnetic response to external 
magnetic fields (red curve). Their saturation magnetization is comparable to ferro or 
ferri-magnetic materials (black curve), but without remanence as in the case of paramagnetic 
materials (blue curve). For a given particle composition, all three behaviors might be 
encountered, depending on temperature or particle size (upper arrows). 
external field. This alignment requires domain wall motions and results in hysteresis 
of the magnetization curve. After the external field is switched off, a remaining 
or remanent magnetization "MR" can be observed. Again, in order to achieve a random 
domain orientation, more energy must be provided by means of an external 
magnetic field applied in the opposite direction. A coercive field "He" is defined as 
the value of the external field necessary to misalign the domains to a random state. 
More detailed information is available in the literature.16-20 
Magnetic materials can be composed of different atoms and ions with various 
magnetic moments. The most well known example is magnetite (Fe304), which consists 
of Fe2+ and Fe3+ ions. The crystallographic structure of such materials determines 
whether or not antif erromagnetic or f errimagnetic properties are present. For 
magnetic drug targeting, only ferro- and ferri-magnetic materials are of interest, as 
they react strongly to external magnetic fields due to their non-zero atomic or lattice 
400 Hiifeli & Chastellain 
unit magnetic moment and the domain structure. Temperature also plays a role in 
the magnitude of magnetic response, as high thermal energy can disturb the atomic 
moment orientation within the domains, leading to paramagnetic behavior. 
When ferro- or ferri-magnetic materials are divided, the obtained nanoparticles 
can become small enough to show single domain structure with a non-zero 
magnetic moment. Depending on the particle size, the thermal energy might be 
high enough to have the particles magnetic moment switch between energetically 
favorable (or easy) directions. These directions are defined by the particle structure, 
especially the crystallography, the shape and the surface. As a result of the moment 
oscillation, the net particle magnetization is zero and no remanent magnetization 
is observed, but the particles still strongly react to external magnetic solicitations 
(see Fig. 1). This behavior, called superparamagnetism, is generally encountered for 
particles that are a few nanometers in size. Superparamagnetism can be influenced 
by magnetic interparticle interactions, which lead to collective behavior of several 
particles acting as one bigger particle. The observation time is also important and 
must be longer than the particle relaxation time, necessary to switch from one to 
the other easy direction. 
2.2. Nanoparticles 
No single definition exists to describe a nanoparticle. Most of the time, an arbitrary 
size range is used ("nanometer sized", from 10~9 to 10~6 m). In view of the 
recent developments in nanotechnology, some people now use the drastic behavior 
changes arising below a critical size (such as the superparamagnetic state described 
earlier) to define nanoparticles. When reducing nanoparticle size, not only does the 
surface over volume ratio increase gradually, but a complete modification of the 
material properties may also occur. This is of primary importance in the biomedical 
field, where a change in size can lead to toxic effects. Many unanswered questions 
remain in this field and legal aspects related to nanoparticles are currently under 
discussion.21 
For use in biomedical applications, ferro-, ferri- or superparamagnetic particles 
must be coated to ensure colloidal stability, increased circulation time in the body, 
functional surfaces, and appropriate diagnostic properties. In this regard, the term 
"magnetic nanoparticle" not only refers to an inorganic core responsible for magnetic 
properties, but also to a composite structure with one or several cores coated 
or embedded in a matrix. Coatings are reviewed elsewhere in this book. 
In addition to a compatible coating, magnetic nanoparticles used in clinical 
applications must form stable aqueous suspensions. Suspensions are complex 
dynamic systems. Their equilibrium is influenced by the forces present, including 
Van der Waals, electrostatic, steric, and magnetic forces, as well as by Brownian 
Magnetic Nanoparticles as Drug Carriers 401 
motion. On this account, it is crucial to realize that solvent modifications can 
drastically influence the behavior of the system. The term "ferrofluid" is correctly 
used only in the case of a colloidal stable suspension of single domain 
nanoparticles.22 
3. Magnetic Nanoparticles 
In general, a single particle type cannot be used for all applications. Instead, the 
composition, size and production route of synthesized magnetic nanoparticles is 
determined by the target application. Although superparamagnetic, f erro- and f errimagnetic 
particles can all be used for magnetic drug carrier applications, superparamagnetic 
particles are favored for biomedical applications, due to the fact that they 
behave non-magnetically when they are not under the influence of an external magnetic 
field, thus preventing undesired magnetic agglomeration. To further assist in 
preventing agglomerations, to optimize bio-interactions with the host environment 
and to maximize biocompatibility, the choice of appropriate surface chemistries 
and functionalizations is also important. Many magnetic nanoparticles are available 
with different surface chemistries, and details about the properties of these 
chemistries are given elsewhere in this book. 
The following subsections provide an overview of magnetic nanoparticles as 
drug carriers, classified according to magnetic composition. The final subsection 
deals with the general biocompatibility issues of magnetic nanoparticles. 
3.1. Iron oxide based magnetic nanoparticles 
In biomedical applications, the most commonly used magnetic nanoparticles are 
superparamagnetic magnetite (Fe304) and maghemite (y-Fe203). This is due to 
their ease of synthesis using chemical or physical approaches,23-33 as well as their 
general bio-compatibility (and FDA approval). Massart's aqueous coprecipitation 
method,34 which leads to particles easily dispersible in water, is the most cited 
method of magnetic nanoparticle preparation. The particle size can be tuned in 
the 3 to 30 nm size range35 and the particles usually show an ellipsoidal shape. 
The stoichiometry ranges from magnetite to maghemite, the two crystallographic 
structures being very similar.35-38 The size distribution is about 10 to 20% [see 
Fig. 2(a)]. Time consuming size sorting procedures allow for further narrowing of 
the size distribution to about 5% in the best case. A thorough characterization of 
such particles was carried out by the group of Jolivet et al.39~i7 
Iron oxide nanoparticles have been synthesized intensively during the past 
decades, but until recently, phase and size control have been problematic. A newly 
developed two-step approach has allowed for much better control over the particle 
402 Hafeli & Chastellain 
Fig. 2. Typical TEM bright field pictures of maghemite nanoparticles. (a) Classical coprecipitation 
synthesis48 and (b) Decomposition at high temperature of organic precursors.49 
Despite its much improved size and shape distribution, the second particle type suffers 
from two major drawbacks for biomedical applications: Biocompatibility and the ease of 
dispersion in water based solvents. 
structure. In this approach, metal particles are first obtained and then oxidized in 
a controlled way.50 Size distributions of better than 5% can be achieved in the 4 to 
16 nm range, as shown by Alivisatos et al.51 and Hyeon et al.49 [see Fig. 2(b)]. These 
particles are, however, often not appropriate for biomedical applications as they do 
not disperse easily in water.52 
Many other magnetite and maghemite nanoparticle synthesis approaches can 
be found in the literature, but none are significantly different from the ones presented 
above. Slightly modified nanoparticles can also be obtained by partly replacing 
the iron in the magnetite or maghemite structure with cobalt or nickel. This in 
turn changes the magnetic properties of the particles. More details are given in a 
recent and extensive review by Tartaj et al.53 
3.2. Cobalt based magnetic nanoparticles 
From a magnetic point of view, particles showing a stronger reaction to magnetic 
fields are desirable. Cobalt achieves this aim, but its toxicity is a major drawback. 
One way of preventing or minimizing this toxicity caused by cobalt ion leakage is 
the inorganic encapsulation of cobalt with, for example, silica. 
3.3. Iron based magnetic particles 
Pure iron nanoparticles can be synthesized, but their sensitivity to oxidation is a 
major drawback for biomedical applications. Thus, a coating, as described for cobalt 
particles, should be used. Iron has also been coated or alloyed with platinum, cobalt 
and carbon. 
Magnetic Nanopartides as Drug Carriers 403 
3.4. Encapsulated magnetic nanopartides 
Depending on the application, magnetic nanopartides may be combined into 
larger conglomerates to increase the overall magnetic moment (see Fig. 3). Great 
care should be paid to interparticle magnetic interactions. The superparamagnetic 
behavior of a system might, for example, be lost due to such interactions. Also, the 
magnetic core concentration must be kept constant among the magnetic conglomerates 
to yield homogeneous magnetic moments, and thus a consistent response to 
an applied magnetic field. 
Either a single or a two-step approach can be used to synthesize magnetic 
particles.54 In the one-step approach, a "linker" is present while synthesizing the 
magnetic nanopartides. In the two-step approach, the linker is added subsequently. 
The linker can be organic or inorganic and is chosen for its chemical and biocompatible 
characteristics. For biomedical applications, dextran, starch, polyethylene 
glycol (PEG), polyvinyl alcohol (PVA), silica and gold are among the most common 
compounds.54 
3.5. Biocompatibility issues of magnetic nanopartides 
One of the first papers to discuss the biocompatibility issues of magnetic particles 
was published in the early 1970's by Nakamura et a/.55 The authors prepared fine 
carbonyl-iron particles and infused them into different animal species in vivo. They 
concluded that to achieve optimal results, the magnetic particles should be coated 
with a biocompatible material and be as round as possible. 
lb) 
f * v , ifps v 
ghemii 
'article 
Fig. 3. Bright field TEM pictures of different types of magnetic particles48: (a) silicamagnetite 
composite and (b) dextran-magnetite composite. The silica layer can be observed 
easily, whereas dextran does not produce enough contrast to be seen clearly. 
404 Hafeli & Chastellain 
Further research showed that pure magnetic metal particles, such as iron, cobalt 
and nickel particles, should not be used directly in vivo because they oxidize easily 
and release +2 or +3 charged metal ions that can exert unwanted as well as 
toxic effects. Iron ions, for example, are problematic in that they produce and catalyze 
oxygen radical formation.56 Cobalt and nickel ions have been found to induce 
adverse tissue reactions, and to promote infection and metal sensitivity.57 
In contrast to pure magnetic particles, iron oxides and superparamagnetic iron 
oxide nanoparticles (SPION) coated and stabilized with hydrophilic polymers have 
been found to be quite thermodynamically stable under physiological conditions, 
not exerting obvious toxic effects. In fact, they are similar in size and core composition 
to the natural non-toxic magnetic nanoparticles found in magnetotactic 
bacteria58 and in human tissue.59 Pharmacokinetic studies of small magnetite 
nanoparticles destined for magnet resonance imaging60,61 have shown that the magnetite 
nanoparticles are taken up by the cells of the reticuloendothelial system (RES) 
and are transported intracellularly to lysosomes, where they slowly oxidize at low 
pH and are then recycled by the body.62 Within 20-40 days, up to 60% of the iron is 
recovered in the red blood cells, as determined using radiolabeled 59Fe. 
Recent discussions have centered on the fate and toxic effects of (magnetic) 
nanoparticles in humans after inhalation. Rodent models have shown the potential 
problematic effects of such particles to include the induction of asthma, inflammation, 
and potentially even cancer.63 Some of these effects might be due to the fact that 
particles smaller than lOOnm are not exhaled, but are almost completely retained 
in the alveoli.64 For this reason, acute effects can rapidly turn into chronic effects. 
Another possible cause for concern is the report that small particles have been found 
behind the blood brain barrier65 (see also Chap. 24). Further research needs to clarify 
if these particles directly crossed the blood brain barrier or via the nose. Care must 
be taken to relate the effects seen in animal models to the human situation, especially 
since effects seen in rodents do not seem to develop in humans.63 Clarification 
of the short and long term risks of nanoparticle use is the aim of several programs 
being initiated in 2005 by the European, American and Canadian governments. 
4. Application of Magnetic Nanoparticles as Drug Carriers 
The following section presents an overview of the use of magnetic nanoparticles 
sized 1 /xm or less for the delivery of drugs. Magnetic microspheres of larger than 
1 /xm size are also mentioned in a few places, but for a recent and thorough review, 
as well as for the history of magnetic drug delivery, the reader is advised to consult 
a more extensive treaty on this matter such as the recent volume of the MML 
series.66 In this section, magnetic nanoparticles will be grouped according to their 
mechanism of action including magnetic hyperthermia; the delivery of magnetic 
nanoparticles that (slowly) release drugs (tumor treatment, thrombolysis, delivery 
Magnetic Nanoparticles as Drug Carriers 405 
of antiinfective, antiarthritic, antifungal, and antiscar agents, and local anesthesia 
or neuroblocking agents) or act without drug release (radiotherapy or embolization); 
and the improved delivery of peptides (gene transfer). Results from in vitro, 
in vivo and clinical work will be discussed. 
4.1. Magnetic hyperthermia 
Magnetic nanoparticles in an alternating current (AC) magnetic field produce heat 
by Neel and Brownian relaxation.67 Heat production above a person's normal body 
temperature is called hyperthermia and can be medically used for the eradication 
of cancer cells. Temperatures above 56°C lead to thermoablation. Once magnetic 
nanoparticles have successfully reached certain organs or tissues, especially tumors, 
magnetic hyperthermia can be induced. Normal tissue nearby, not containing the 
magnetic nanoparticles, remains at body temperature and is thus spared. 
One of the first to examine this effect was Gilchrist who published a seminal 
paper in 1956 on the selective inductive heating of lymph nodes, after injection of 
maghemite particles sized between 20 and 100 nm diameter directly into the lymph 
nodes near surgically removed canine tumors.68 Using 5 mg of Fe203 per gram 
of lymph node and a magnetic field strength of 240 Oe, a maximum temperature 
rise of 14°C was reached within 3 minutes. To prevent a reoccurrence of the cancer, 
hyperthermia is normally combined with a second treatment modality such as 
chemotherapy or irradiation. 
Twenty years later, Rand et al. showed that ferrosilicone can induce heat after 
being infused into a tumor's blood supply and placed under the influence of a strong 
magnetic alternating field.69 Rand's ferromagnetic silicone microspheres were 
based on Turner et al.'s research in 1973, in which magnetic particles of unknown 
but probably larger size in a silicone fluid were infused into and then clogged 
(embolized) the capillary bed of several targeted organs.70 With this technique, the 
researchers successfully embolized the blood supplies of different tumors. No side 
effects were reported in the 7 patients who had brain tumors, pheochromocytomas, 
a tongue tumor and a hypernephroma. Rand's so-called "magnetic field induced 
hyperthermia" was then further developed by Sako et al.,71,72 who showed that 
heating was reproducible and proportional to the amount of iron used. 
The use of single domain, dextran-coated magnetite nanoparticles for tumor 
hyperthermia was developed by Jordan73 and Chan,74 and is currently undergoing 
in vivo and clinical testing. Jordan reported the optimal nanoparticle core diameter to 
be in the 10 nm range,73,75 although type of coating and coating stability also seemed 
to be important.76 Using these nanoparticles, 5 mg of material per gram of tumor 
was sufficient to increase the tumor temperature by 10°C to cell toxic levels. Jordan is 
currently conducting a clinical phase II trial combining magnetic hyperthermia and 
radiation therapy77'78 He recently presented the first clinical results from 8 patients 
406 Hafeli & Chastellain 
at the 5th International Conference on the Scientific and Clinical Applications of 
Magnetic Carriers in Lyon, France. The patients were treated for cervix (2), rectal, 
and prostate (2) carcinoma, as well as for a chondrosarcoma, rhabdomyosarcoma, 
and liver metastasis. During the 1-hour sessions, after local injection of the magnetic 
particles, the tumor temperature increased to 43-50° C under the influence of a 
magnetic field of 3 to 9.5kA/m and a frequency of 100 kHz. While no additional 
nanoparticle injections were necessary, the hyperthermia treatment was repeated 
from 2-11 times. The magnetic fluid hyperthermia was well tolerated. Two patients 
showed complete remission 9 and 14 months after treatment, while the other six 
patients showed local control with no recurrent growth of the tumors. These results 
are very promising. 
Another group in Germany led by Hilger is working on circumventing the 
drawback of having to directly inject the particles into a tumor, by using antibodybound 
magnetic nanoparticles which are able to target breast cancer, followed by 
magnetic field hyperthermia.79'80 Although their particles are taken up extensively 
by tumor cells and show a specific heating power of up to 170 W/g,81 there is still 
more work needed to increase the number of particles in the tumor and to reach a 
homogeneous tumor distribution.82 
Magnetic hyperthermia is also possible with large microspheres that contain 
magnetic nanoparticles.83,84 As an example, Moroz et al. incorporated 100 nm 
maghemite particles into 32 /xm biocompatible plastic particles and then embolized 
the arterial blood supply of liver tumors with them. In an animal study with 10 rabbits, 
the VX2 tumor volumes decreased significantly within 2 weeks.85-87 
The development of maghemite nanoparticles with very high AC losses is ongoing. 
Hergt et al. are in the process of characterizing the largest, but still superparamagnetic 
particles,88 optimizing the coatings such as carboxydextran or polyethylene 
glycol,89 and investigating the exact mechanism of heat production in an AC magnetic 
field.88 
Magnetic hyperthermia is an exciting cancer treatment possibility and is profiting 
from ongoing research into its mechanism of action and from improved magnetic 
materials. The proof of principle has advanced to the clinical stage with the 
construction and clinical testing of Jordan's magnetic field therapy system.77 The 
targeted (cancer) cell uptake of sufficient amounts of magnetic nanoparticles from 
a patient's blood supply could make magnetic hyperthermia the method of choice 
for many different kinds of tumors. 
4.2. Magnetic chemotherapy 
Magnetic drug delivery is able to concentrate drugs in a tumor if the tumor is 
accessible through the arterial system and has a good supply of blood. Magnetic 
Magnetic Nanoparticles as Drug Carriers 407 
drug delivery thus promises to deliver highly effective anticancer drugs with fewer 
side effects, and with shorter and less toxic treatments. 
Most drug release from magnetic particles occurs passively, by desorption from 
and diffusion out of the particle matrix. The main driving forces are pH, osmolarity 
and concentration differences between particles and the blood/tissue. Widder and 
Senyei were the first to successfully illustrate this concept with the chemotherapeutic 
drug doxorubicin encapsulated into albumin-coated magnetite particles sized 
around 1-2 /xm.90 Targeting a distinct area of a rat's tail, they were able to deliver 
200 times more of the drug than intravenous application of the same amount of 
free drug could achieve.91 Taking it a step further, they treated Yoshida rats with 
sarcomas in their tails and attained complete remission in 77% of the rats.92,93 
The magnetic albumin microspheres were never tested clinically, likely because 
the magnetophoretic mobility (overall magnetic responsiveness to a magnetic field) 
was considered too low for deeper applications. This changed with the introduction 
of iron-carbon particles originally developed in Russia94 and then brought to clinical 
trial by the company FeRx. FeRx's irregularly-shaped carbon-coated iron particles, 
of 0.5 to 5 /xm in diameter and with a very high magnetic susceptibility, were loaded 
with doxorubicin and showed promising results and very low therapy-related toxicity 
in the treatment of inoperable liver cancer.95'96 Unfortunately, FeRx ceased to 
exist in 2004 when a preliminary analysis of their ongoing clinical trial failed to 
convince investors of the method's superiority over other treatment methods. 
Not only doxorubicin, but also many other chemotherapeutic drugs can be 
and have been adsorbed to magnetic nanoparticles made from many different 
matrix materials. Examples of chemotherapeutic magnetic nanoparticles tested 
in vivo include polyalkylcyanoacrylate nanoparticles of 220 nm diameter filled 
with (adsorbed) dactinomycin,97 chitosan nanoparticles of 530 nm diameter loaded 
with oxantrazole,98 solid lipid nanoparticles of 450-570 nm diameter loaded with 
methotrexate,99 and ferro carbon of 100 nm diameter loaded with carminomycine.94 
Each of these nanoparticles has been tested in animal experiments with positive 
results. Specifically, after intravenous injection, the drug concentration tripled in 
the target organ when a magnet was placed above it, compared with a control 
without an applied magnet. 
It seems that the intravenous injection of magnetic nanoparticles, even very 
close to the target region, is not optimal. This was well documented in a clinical 
cancer therapy trial performed by Liibbe et al. in 14 patients.100,101 They used 
magnetic nanoparticles of 100 nm in diameter loaded with 4'-epidoxorubicin for 
the treatment of advanced solid cancer. The phase I study clearly showed accumulation 
of magnetic nanoparticles in the target area without toxic effects. MRI 
measurements, however, indicated that more than 50% of the magnetic nanoparticles 
were deposited in the liver. This was likely due to the particles' low magnetic 
408 Hafeli & Chastellain 
susceptibility and small size, which limited their ability to be held at the target 
organ. Intraarterial injection into the blood supply that leads to the target region 
might be much more effective for magnetic drug targeting for this reason. 
The above examples of magnetic drug targeting with magnetic nanospheres 
are only a subset of all the magnetic drug delivery attempts. A more complete compilation 
is given by Hafeli.66 In addition, all the important factors in magnetically 
controlled targeted chemotherapy are extensively described in a review by Gupta 
and Hung.102 
4.3. Other magnetic treatment approaches 
Under the influence of a magnetic field, magnetic particles align in chains and 
eventually agglomerate. Depending on particle size and shape, this can lead to 
embolization (clogging) of the blood vessels and especially of the small capillaries 
of 7 to 10 ^m in diameter. This accumulation of particles can be used on its own 
to starve the target tissue of oxygen, produce hypoxia and induce necrosis. The 
magnetic particles used for this approach are generally larger, such as the "iron 
sponge" of 10-30 fim used by Sako et a/.,103,104 but can also consist of nanoparticles in 
a more lipophilic solvent such as the ferrosilicone employed by Turner et al.70 Turner 
added a catalyst to their ferrosilicone suspension, which resulted in vulcanization of 
the viscous slurry 14 min after injection into 7 patients with diverse solid tumors.70 
Magnetic particles can also be used for tumor treatment without releasing any 
drugs. For this purpose, magnetic particles can incorporate radioisotopes either in 
the matrix or bound to their surface, and then deliver tumor cell-toxic radiation 
doses wherever they accumulate.105 External magnetic fields or internal magnetizable 
wires106,107 can be used to accumulate radioactive magnetic particles and 
hold them at the target site. The particles irradiate the area within the specific treatment 
range of the isotope. Initial experiments in mice showed that intraperitoneally 
injected radioactive polydactic acid) based magnetic microspheres (10-20/zm in 
diameter) could be concentrated near a subcutaneous tumor in the belly area above 
which a small magnet had been attached.108 The dose-dependent irradiation from 
the /J-emitter 90Y-containing magnetic particles resulted in the complete disappearance 
of more than half of the tumors. 
Iron carbon-based smaller radioactive particles of around 1/xm have been 
radiolabeled with different diagnostic (99mTc, mIn) and therapeutic (188Re, 90Y) 
radioisotopes.109,110 Targeting studies to distinct liver regions in swine by our group 
(unpublished results) showed that more than 90% of the injected radioactive magnetic 
particles were accumulated underneath a strong NdFeB-magnet. The radioactive 
particles stayed in the target region for at least 3 days, even after the removal 
of the magnet. 
Magnetic Nanoparticles as Drug Carriers 409 
Magnetic nanoparticles of much smaller diameters are being used clinically 
for diagnostic purposes, mainly as contrast agents.111 The accumulation of these 
magnetic particles is, however, based on non-specific properties such as the tissuespecific 
pore size (fenestration) or enhanced permeability and retention effect (EPR) 
seen in tumor tissues, but not on magnetic targeting. Recent examples of nonspecific 
targeting are the internalization into cells of the positively charged at 
peptide bound to therapeutic agents, such as radioactively complexed 99mTc and 
188Re,112 and superparamagnetic iron oxides known as tat-CLIO (tat-cross linked 
iron oxides).113 
Magnetic drug delivery is also able to deliver other types of drugs such as highly 
potent antiinfective, blood clot-dissolving, anti-inflammatory, anti-arthritic, photodynamic 
therapy and paralysis-inducing drugs, among many others.66 A good 
example of these applications was reported in 1988.114 In this study, Torchilin et al. 
surgically induced a thrombus in both carotid arteries of a dog, fixed a permanent 
magnet near one of them, and 1 hr later, intravenously injected 1 /xm-sized dextran 
microspheres with covalently bound streptokinase. The side without the magnet 
completely occluded within 4 hrs, while the magnet side returned to initial blood 
flow conditions after about 30 min and appeared completely open at histological 
examination. Torchilin noted that these results were achieved using doses 10 times 
lower than those used when streptokinase is directly injected. Similar results were 
obtained in the same year by another group using 30-60 nm PEGylated magnetite 
particles containing urokinase.115 In both cases, the thrombolytic activity remained 
at background levels outside the targeted region. 
4.4. Magnetic gene transfer 
The newest application of magnetic nanoparticles is for targeted and enhanced 
gene delivery in potential applications such as wound repair116 and the treatment of 
cancer,117,118 eye disease,119 and cystic fibrosis.120 Magnetically enhanced gene transfer 
may be able to overcome the current lack of selectivity of the existing vectors and 
low efficiency of gene transfer. The mechanisms by which magnetic nanoparticles 
can improve on transfection rates are by magnetically forced contact121'122 and by 
increasing the plasmid concentration magnetically.123 
Magnetofection was first described in a Japanese patent by Harata et al. who 
used magnetic liposomes to transfect cells.124 Bergemann et al. described the first 
experiments of transfecting cytokine-induced killer cells (CIK-cells) with plasmid 
DNA carrying distinct interleukin genes. However, their magnetic nanoparticles 
were only used as plasmid carriers, not as the driving force, which was provided by 
electroporation.12 Using a magnet as the driving force was then described a couple 
of years later by Plank et al.122,125 For successful in vitro and in vivo transfections, 
410 Hafeli & Chastellain 
only very small amounts of plasmid were necessary, and the transfection occurred 
in a matter of just a couple of minutes. This speedy and efficient transfection at 
low vector doses is the main advantage of magnetofection. Furthermore, remotely 
controlled vector targeting in vivo seems possible. 
5. Conclusions 
The technological advancements in the material and engineering sciences, and especially 
in the nanotechnology revolution, with its increasing molecular approach 
to the synthesis, derivatization, combination, self-assembly, and manipulation of 
materials, will guide improvements in all aspects of magnetic nano- and microparticles. 
These advancements include the synthesis of higher magnetic nanophases; 
the increasing availability of stronger magnets and engineered magnetic fields; the 
ability to prepare more uniform particles; and the rendering of these particles biocompatible 
and ultimately biodegradable without toxicity. 
For real therapeutic breakthrough, however, a few challenges in the field of 
magnetic drug delivery still need to be addressed. One of the challenges is the difficulty 
involved in generating the focused field profiles needed to target magnetic 
nanoparticles deep within the body, due to the speed with which the magnetic fields 
drop off as their distance from the source increases (intensity = 1 / r3). Another challenge 
centers on attaining homogeneous particle distributions given that blood flow 
in the target region can vary from very fast (100-200 cm/s) to very slow (0.05 cm/s). 
A final challenge is the optimization of magnetic nanoparticles in terms of magnetization 
and size uniformity. All these problematic areas are currently being 
addressed by multidisciplinary groups worldwide, as evident from a special issue 
of the / Magn Magn Mater 293:1-736, containing 107 original peer-reviewed papers 
that were submitted at the 5th International Conference on the Scientific and Clinical 
Applications of Magnetic Carriers in 2004 (www.magneticmicrosphere.com). 
The potential of magnetic drug delivery is great. In addition to their magnetic 
responsiveness, magnetic nanoparticles carry an innate signal that can be used 
for magnetic resonance imaging.126 Furthermore, other imaging modalities such 
as radioisotope or fluorescence imaging can be used after derivatizing the particle 
surface. There is no limitation to the kind of drugs that can be encapsulated or 
bound to magnetic nanoparticles.66 Also, current pharmaceutical techniques allow 
for the development of drug release profiles for a large group of drugs and diseases. 
Magnetic targeting devices such as the recently FDA approved Niobe system 
(Stereotaxis Inc., St. Louis, Missouri, U.S.A.)127 improve on the precise manipulation 
of magnetic forces. Anatomical and physiological conditions in a patient are, 
however, complicated, and successful therapies will have to be specifically adapted 
to each disease and ideally applied under imaging control. 
Magnetic Nanoparticles as Drug Carriers 411 
The treatment of cancer is an especially good target for magnetic d r u g delivery. 
However, any disease that could benefit from precise control over the delivery of 
highly effective, but potentially toxic substances would also be a good candidate. 
For example, antiangionic drugs could be delivered to the back of the eye to prevent 
blindness in patients with age-related macular degeneration. Attempts to develop 
such a drug delivery system are ongoing.128,129 
References 
1. Massover WH (1993) infrastructure of ferritin and apoferritin: A review. Micron 
24:389^137. 
2. Chasteen ND and Harrison PM (1999) Mineralization of ferritin: An efficient means of 
iron storage. / Struct Biol 126:182-194. 
3. Hafeli UO (2006) The history of magnetism in medicine, in Andra W and Nowak H 
(eds.), Magnetism in Medicine: A Handbook, 2nd edn. Wiley-VCH: Berlin, 1-25. 
4. Shinkai M (2002) Functional magnetic particles for medical applications. / Biosci Bioeng 
94:606-613. 
5. Roger J, Pons JN, Massart R, Halbreich A and Bacri JC (1999) Some biomedical applications 
of ferrofluids. Eur Phys ] Appl Phys 5:321-325. 
6. Liibbe AS, Bergemann C, Brock J and McClure DG (1999) Physiological aspects in 
magnetic drug targeting. } Magn Magn Mater 194:149-155. 
7. Taylor JI, Hurst CD, Davies MJ, Sachsinger N and Bruce IJ (2000) Application of magnetite 
and silica-magnetite composites to the isolation of genomic DNA. / Chromatography 
A 890:159-166. 
8. Jung CW and Jacobs P (1995) Physical and chemical properties of superparamagnetic 
iron oxide MR contrast agents: Ferumoxides, ferumoxtran, ferumoxsil. J Magn Res Imag 
13:661-674. 
9. Jung CW, Rogers JM and Groman EV (1999) Lymphatic mapping and sentinel node 
location with magnetite nanoparticles. / Magn Magn Mater 194:210-216. 
10. Safarik I and Safarikova M (1999) Use of magnetic techniques for isolation of cells. 
/ Chromatography B 772:33-53. 
11. Goodwin S, Peterson C, Hoh C and Bittner C (1999) Targeting and retention of magnetic 
targeted carriers (MTCs) enhancing intra-arterial chemotherapy. / Magn Magn Mater 
194:132-139. 
12. Bergemann C, Muller-Schulte D, Oster J, Brassard A and Liibbe AS (1999) Magnetic ionexchange 
nano- and microparticles for medical, biochemical and molecular biological 
applications. / Magn Magn Mater 194:45-52. 
13. Sheng R, Flores GA and Liu J (1999) In vitro investigation of a novel cancer therapeutic 
method using embolizing properties of magnetorheological fluids.} Magn Magn Mater 
194:167-175. 
14. Alexiou C et ah (2006) Targeting cancer cells: Magnetic nanoparticles as drug carriers. 
Eur Biophys J 35:446-450. 
412 Hafeli & Chastella in 
15. Richert H et al. (2005) Development of a magnetic capsule as a drug release system for 
future applications in the human GI tract. J Magn Magn Mater 293:497-500. 
16. Chen CW (1986) Magnetism and Metallurgy of Soft Magnetic Materials. Dover Publications, 
Amsterdam. 
17. Tebble RS and Craik DJ (1969) Magnetic Materials. Wiley Interscience, New York. 
18. Cullity BD and Graham CD, (1972) Introduction to Magnetic Materials. Wiley, New York. 
19. Dormann JL (1997) Magnetic relaxation in fine particles systems, in Prigogine I and 
Rice SA (eds.), Adv Chem Phys, Vol. 98, Wiley, New York. 
20. BatUe X and Labarta A (2002) Finite-size effects in fine particles: Magnetic and transport 
properties. J Phys D 35:R15-R42. 
21. Bainbridge WS (2004) Social and ethical implications of nanotechnology, in Bhushan B, 
(ed.) Handbook of Nanotechnology. Springer: Berlin, 1135-1151. 
22. Charles SW (2001) Magnetic fluids (ferrofluids), in Dormann JL and Fiorani D (eds.) 
Studies of Magnetic Properties of Fine Particles. Elsevier Science Publishers: Amsterdam, 
267-276. 
23. Rosensweig RE (1982) Magnetic fluids. Sci Am 247(4):136-145. 
24. Zdujic M et al. (1999) The ball milling induced transformation of alpha-Fe2C>3 powder 
in air and oxygen atmosphere. Mater Sci Eng A 262:204-213. 
25. Ozaki M and Matijevic E (1985) Preparation and magnetic properties of monodispersed 
spindle-type y-Fe2C>3 particles. / Colloid Interf Sci 107:199. 
26. Itoh H and Sugimoto T (2002) Systematic control of size, shape, structure, and magnetic 
properties of uniform magnetite and maghemite particles. / Colloid Interf Sci 
265:283-295. 
27. Ozaki M (1989) Preparation and properties of well-defined magnetic particles. MRS 
BuZZ 14:41. 
28. Bica I (2003) On the mechanisms of iron microspheres formation in argon plasma jet. 
fMagn Magn Mater 257:119-125. 
29. Yitai Q et al. (1994) Hydrothermal preparation and characterization of ultrafine magnetite 
powders. Mater Res Bull 29:953-957. 
30. Bae D-S, Han K-S, Cho S-B and Choi S-H (1998) Synthesis of ultrafine Fe304 powder 
by glycothermal process. Mater Lett 37:255-258. 
31. Bomati-Miguel Oetal. (2005) Fe-based nanoparticulate metallic alloys as contrast agents 
for magnetic resonance imaging. Biomaterials 26:5695-5703. 
32. Vijayakumar R, Koltypin Y, Felner I and Gedanken A (2000) Sonochemical synthesis 
and characterization of pure nanometer-sized Fe304 particles. Mater Sci 
Eng, A 286:101-105. 
33. Daichuan D, Pinjie H and Shushan D (1995) Preparation of uniform a-Fe203 particles 
by microwave-induced hydrolysis of ferric salts. Mater Res Bull 30:531-541. 
34. Massart R (1980) Magnetic fluids and process for obtaining them. U. S. Patent 
No. 4,329,241 (May 11). 
35. Massart R and Cabuil V (1987) Synthese en milieu alcalin de magnetite colloidale. 
/ Chim Phys 84:963-973. 
36. Cornell RM and Schwertmann U (1996) The Iron Oxides: Structure, Properties, Reactions, 
Occurrences and Uses. VCH: Weinheim. 
Magnetic Nanoparticles as Drug Carriers 413 
37. Nasrazadani S and Raman A (1993) The application of infrared spectroscopy to the 
study of rust systems — II. Study of cation deficiency in magnetite (Fe304) produced 
during its transformation to maghemite (y-Fe2C>3) and hematite (a-Fe2C>3). Corros Sci 
34:1355-1365. 
38. Kaczmarek WA (1996) Structural and magnetic properties of cobalt-doped iron 
oxide particles prepared by novel mechanochemical method. / Magn Magn Mater 
157/158:264-265. 
39. Tronc E and Jolivet JP (1986) Surface effects on magnetically coupled gamma-Fe203 
colloids. Hyperfine Interact 28:525-528. 
40. Fiorani D, Testa AM, Lucari F, D'Orazio F and Romero H (2002) Magnetic properties 
of maghemite nanoparticle systems: Surface anisotropy and interparticle interaction 
effects. Physica B (Amsterdam) 320:122-126. 
41. Tronc E, Ezzir A, Cherkaoui R, Chaneac C, Nogues M, Kachkachi H, Fiorani D and 
Jolivet J-P (2000) Surface-related properties of gamma-Fe2C>3 nanoparticles. / Magn 
Magn Mater 221:63-79. 
42. Dormann JL, Cherkaoui R, Spinu L, Nogues M, Lucari F, D'Orazio F, Fiorani D and 
Jolivet J-P (1998) From pure superparamagnetic regime to glass collective state of magnetic 
moments in y-Fe203 nanoparticle assemblies.} Magn Magn Mater 187:L139-L144. 
43. Gazeau F, Bacri JC, Gendron F, Perzynski R, Raikher YL, Stepanov VI and Dubois E 
(1998) Magnetic resonance of ferrite nanoparticles: Evidence of surface effects. / Magn 
Magn Mater 186:175-187. 
44. Fiorani D, Dormann JL, Cherkaoui R, Tronc E, Lucari F, D'Orazio F, Spinu L, Nogues M, 
Garcia Aand Testa AM (1999) Collective magnetic state in nanoparticles systems. J Magn 
Magn Mater 196:143-147. 
45. Dormann JL, Dormann L, Spinu L, Tronc E, Jolivet JP, Lucari F, D'Orazio F and 
Fiorani D (1998) Effect of interparticle interactions on the dynamical properties of y- 
Fe203 nanoparticles. / Magn Magn Mater 183:L255-L260. 
46. Prodan D, Chaneac C, Tronc E, Jolivet J-P, Cherkaour R, Ezzir A, Nogues M and 
Dormann JL (1999) Adsorption phenomena and magnetic properties of y-Fe203 
nanoparticles. / Magn Magn Mater 203:63-65. 
47. Tronc E, Fiorani D, Nogues M, Testa AM, Lucari F, D'Orazio F, Greneche JM, 
Wernsdorfer W, Galvez N, Chaneac C, Mailly D and Jolivet JP (2003) Surface effects in 
noninteracting and interacting y-Fe203 nanoparticles. / Magn Magn Mater 262:6-14. 
48. Chastellain M (2004). Nanoscale superparamagnetic composite particles for biomedical 
applications. Ph.D. thesis, Ecole Polytechnique Federale, Lausanne. 
49. Hyeon T, Su Seong Lee, Park J, Chung Y and Hyon Bin Na (2001) Synthesis of highly 
crystalline and monodisperse maghemite nanocrystallites without a size-selection process. 
JAmer Chem Soc 123:12798-12801. 
50. Murray CB, Sun S, Doyle H and Betley T (2001) Monodisperse 3d Transition-Metal 
(Co, Ni, Fe) nanoparticles and their assembly into nanoparticle superlattices. MRS Bull 
26:985. 
51. Rockenberger J, Scher EC and Alivisatos AP (1999) A new nonhydrolytic singleprecursor 
approach to surfactant-capped nanocrystals of transition metal oxides.} Am 
Chem Soc 121:11595-11596. 
414 Hafeli & Chastellain 
52. Charles SW (2001) Ferrofluids: Preparation and physical properties, in Buschow KHJ 
et al. (eds.) Encyclopedia of Materials: Science and Technology. Elsevier: Amsterdam. 
53. Tartaj P, del Puerto Morales M, Veintemillas-Verdaguer S, Gonzalez-Carreno T 
and Serna CJ (2003) The preparation of magnetic nanoparticles for applications in 
biomedicine. / Phys D Appl Phys 36:R182-R197. 
54. Arshady R (2001) Microspheres, Microcapsules & Liposomes: Magneto- and Radiopharmaceuticals, 
1st edn. Citus Books, London. 
55. Nakamura T, Konno K, Morone T, Tsuya N and Hatano M (1971) Magneto-medicine: 
Biological aspects of ferromagnetic fine particles. / Appl Physiol 42:1320-1324. 
56. Darley-Usmar V and Halliwell B (1996) Blood radicals: Reactive nitrogen species, 
reactive oxygen species, transition metal ions, and the vascular system. Pharm Res 
13:649-662. 
57. Granchi D et al. (1998) Cell death induced by metal ions: Necrosis or apoptosis? } Mater 
Sci Mater Med 9:31-37. 
58. Schiiler D (1999) Formation of magnetosomes in magnetotactic bacteria. JMol Microbiol 
Biotechnol 1:79-86. 
59. Kirschvink JL, Kobayashi-Kirschvink A, Diaz-Ricci JC and Kirschvink SJ (1992) Magnetite 
in human tissues: A mechanism for the biological effects of weak ELF magnetic 
fields. Bioelectromagnetics Suppl. 1:101-113. 
60. Van Hecke P, Marchal G, Decrop E and Baert AL (1989) Experimental study of the 
pharmacokinetics and dose response of ferrite particles used as contrast agent in MRI 
of the normal liver of the rabbit. Invest Radiol 24:397-399. 
61. Weissleder R et al. (1989) Superparamagnetic iron oxide: Pharmacokinetics and toxicity. 
Am J Roentgenol 152:167-173. 
62. Lawaczeck R et al. (1997) Magnetic iron oxide particles coated with carboxydextran for 
parenteral administration and liver contrasting. Pre-clinical profile of SH U555A. Acta 
Radiol 38:584-597. 
63. Warheit DB (2004) Nanoparticles: Health impacts? Materials Today February, 32-35. 
64. ICRP Report 66 (1994) Human respiratory tract model for radiological protection. 
International Commission on Radiological Protection, Oxford. 
65. Kreuter J (2004) Influence of the surface properties on nanoparticle-mediated transport 
of drugs to the brain. / Nanosci Nanotechnol 4:484-488. 
66. Hafeli UO (2006) Magnetic nano- and microparticles for targeted drug delivery, in 
Arshady R and Kono K (eds.), Smart nanoparticles in nanomedicine — the MML series, 
Vol. 8. Kentus Books, London, UK, pp. 77-126. 
67. Hergt R, Hiergeist R, Hilger I and Kaiser W (2002) Magnetic nanoparticles for thermoablation, 
in Pandalai SG (ed.), Recent Research Developments in Materials Science, Vol. 3 
Part 2, pp. 723-742. 
68. Gilchrist RK et al. (1957) Selective inductive heating of lymph nodes. Ann Surg 
146:596-606. 
69. Rand RW, Snyder M, Elliott D and Snow H (1976) Selective radiofrequency heating 
of ferrosilicone occluded tissue: A preliminary report. Bull Los Angeles Neurol Soc 
41:154-159. 
Magnetic Nanoparticles as Drug Carriers 415 
70. Turner RD, Rand RW, Bentson JR and Mosso JA (1975) Ferromagnetic silicone necrosis 
of hypernephromas by selective vascular occlusion to the tumor: A new technique. 
/ Urol 113:455^59. 
71. Sako M and Hirota S (1986) Embolotherapy of hepatomas using ferromagnetic microspheres, 
its clinical evaluation and the prospect of its use as a vehicle in chemoembolohyperthermic 
therapy. Gan To Kagaku Ryoho 13:1618-1624. 
72. Hase M et al. (1989) Experimental study of embolo-hyperthermia for treatment of liver 
tumor-induction heating to ferromagnetic particles injected into tumor tissue. Nippon 
Igaku Hoshasen Gakkai Zasshi 49:1171-1173. 
73. Jordan A et al. (1993) Inductive heating of ferrimagnetic particles and magnetic fluids: 
Physical evaluation of their potential for hyperthermia. Int J Hyperthermia 9:51-68. 
74. Chan DCF, Kirpotin DB and Bunn PA (1993) Synthesis and evaluation of colloidal magnetic 
iron oxides for the site-specific radiofrequency-induced hyperthermia of cancer. 
JMagn Magn Mater 122:374-378. 
75. Jordan A, Rheinlander T, Waldofner N and Scholz R (2003) Increase of the specific 
absorption rate (SAR) by magnetic fractionation of magnetic fluids. / Nanoparticle Res 
5:597-600. 
76. Jordan A et al. (1996) Cellular uptake of magnetic fluid particles and their effects 
in AC magnetic fields on human adenocarcinoma cells in vitro. Int } Hyperthermia 
12:705-722. 
77. Jordan A et al. (2001) Presentation of a new magnetic field therapy system for the 
treatment of human solid tumors with magnetic fluid hyperthermia. / Magn Magn 
Mater 225:118-126. 
78. Jordan A. (2003) Magnetized iron particles melt tumors. Arztezeitung. 
79. Hilger I et al. (2001) Electromagnetic heating of breast tumors in interventional radiology: 
In vitro and in vivo studies in human cadavers and mice. Radiology 218:570-575. 
80. Hilger I et al. (2002) Heating potential of iron oxides for therapeutic purposes in interventional 
radiology. Acad Radiol 9:198-202. 
81. Hilger I et al. (2004) Magnetic nanoparticles for selective heating of magnetically labeled 
cells in culture: Preliminary investigation. Nanotechnology 15:1027-1032. 
82. Hilger letal. (2002) Thermal ablation of tumors using magnetic nanoparticles: An in vivo 
feasibility study. Invest Radiol 37:580-586. 
83. Minamimura T et al. (2000) Tumor regression by inductive hyperthermia combined 
with hepatic embolization using dextran magnetite-incorporated microspheres in rats. 
Int J Oncol 16:1153-1158. 
84. Moroz P, Jones SK and Gray BN (2002) Magnetically mediated hyperthermia: Current 
status and future directions. Int ] Hyperthermia 18:267-284. 
85. Moroz P, Jones SK, Winter J and Gray BN (2001) Targeting liver tumors with hyperthermia: 
Ferromagnetic embolization in a rabbit liver tumor model. / Surg Oncol 
78:22-29. 
86. Moroz P, Jones SK and Gray BN (2002) Tumor response to arterial embolization hyperthermia 
and direct injection hyperthermia in a rabbit liver tumor model. / Surg Oncol 
80:149-156. 
41 6 Hafeli & Chastellain 
87. Moroz P, Jones SK and Gray BN (2002) The effect of tumour size on ferromagnetic 
embolization hyperthermia in a rabbit liver tumour model. Int } Hyperthermia 
18:129-140. 
88. Hergt R et al. (2004) Enhancement of AC-losses of magnetic nanoparticles for heating 
applications. / Magn Magn Mater 280:358-368. 
89. Hergt R et al. (2004) Maghemite nanoparticles with very high AC-losses for application 
in RF-magnetic hyperthermia. / Magn Magn Mater 270:345-357. 
90. Widder K, Flouret G and Senyei A (1979) Magnetic microspheres: Synthesis of a novel 
parenteral drug carrier. / Pharm Sci 68:79-82. 
91. Senyei AE, Reich SD, Gonczy C and Widder KJ (1981) In vivo kinetics of magnetically 
targeted low-dose doxorubicin. / Pharm Sci 70:39-41. 
92. Widder KJ, Morris RM, Howard DP and Senyei AE (1981) Tumor remission in Yoshida 
sarcoma-bearing rats by selective targeting of magnetic albumin microspheres containing 
doxorubicin. Proc Natl Acad Sci USA 78:579-581. 
93. Widder KJ, Morris RM, Poore GA, Howards DP and Senyei AE (1983) Selective targeting 
of magnetic albumin microspheres containing low-dose doxorubicin: Total remission 
in Yoshida sarcoma-bearing rats. Eur J Cancer Clin Oncol 19:135-139. 
94. Kuznetsov AA et al. (1997) Ferro-carbon particles: Preparation and applications, in 
Hafeli U et al. (eds.), Scientific and Clinical Applications of Magnetic Carriers. Plenum 
Press, New York, pp. 379-389. 
95. Goodwin S (2000) Magnetic targeted carriers offer site-specific drug delivery. Oncol 
News Int 9:22. 
96. Johnson J et al. (2002) The MTC technology: A platform technology for the site-specific 
delivery of pharmaceutical agents. Eur Cells Mat 3:12-15. 
97. Ibrahim A, Couvreur P, Roland M and Speiser P (1983) New magnetic drug carrier. 
/ Pharm Pharmacol 35:59-61. 
98. Hassan EE and Gallo JM (1993) Targeting anticancer drugs to the brain, I: Enhanced 
brain delivery of oxantrazole following administration in magnetic cationic microspheres. 
/ Drug Targ 1:7-14. 
99. Vyas SP and Jaitely V (1999) Magnetoresponsive solid lipid nanoparticles (SLN) as 
novel targeting modules for targeting of methotrexate. Proceed Intern Symp Control Rel 
Bioact Mater 26, Abstract #6237. 
100. Liibbe AS et al. (1996) Clinical experiences with magnetic drug targeting: A phase I 
study with 4'-epidoxorubicin in 14 patients with advanced solid tumors. Cancer Res 
56:4686-4693. 
101. Liibbe AS, Alexiou C and Bergemann C (2001) Clinical applications of magnetic drug 
targeting. / Surg Res 95:200-206. 
102. Gupta PK and Hung CT (1994) Magnetically controlled targeted chemotherapy, in 
Willmott N and Daly J (eds.), Microspheres and Regional Cancer Therapy. CRC Press, Boca 
Raton, pp. 71-116. 
103. Sako M et al. (1982) Transcatheter microembolization with ferropolysaccharide: A new 
approach to ferromagnetic embolization of tumors: Preliminary report. Invest Radiol 
17:573-582. 
Magnetic Nanoparticles as Drug Carriers 417 
104. Sako M, Hirota S and Ohtsuki S (1985) Clinical evaluation of ferromagnetic 
microembolization for the treatment of hepatocellular carcinoma. Ann Radiol 
29:200-204. 
105. Hafeli UO (2001) Radiolabeled magnetic microcapsules for magnetically targeted 
radionuclide therapy, in Arshady R (ed.), Microspheres, Microcapsules & Liposomes: Radiolabeled 
and Magnetic Particulates in Medicine and Biology, Vol. 3. Citus Books, London, 
pp. 559-584. 
106. lacob G, Rotariu O, Strachan NJC and Hafeli UO (2004) Magnetizable needles and 
wires-modeling an efficient way to target magnetic microspheres in vivo. Biorheology 
41:599-612. 
107. Rotariu O, lacob G, Strachan NJC and Chiriac H (2004) Simulating the embolization 
of blood vessels using magnetic microparticles and acupuncture needle in a magnetic 
field. Biotechnol Prog 20:299-305. 
108. Hafeli UO, Pauer GJ, Roberts WK, Humm JL and Macklis RM (1997) Magnetically 
targeted microspheres for intracavitary and intraspinal Y-90 radiotherapy, in Hafeli U, 
Schiitt W, Teller J and Zborowski M (eds.), Scientific and Clinical Applications of Magnetic 
Carriers, 1st edn. Plenum, New York, 501-516. 
109. Yu JF et al. (2002) Radiolabeling of magnetic targeted carriers with several therapeutic 
and imaging radioisotopes. Eur Cells Mat 3(Suppl. 2):16-18. 
110. Hafeli UO, Yu J, Farudi F, Li Y and Tapolsky G (2003) Radiolabeling of magnetic targeted 
carriers (MTC) with indium-Ill. Nucl Med Biol 30:761-769. 
111. Wang YX, Hussain SM and Krestin GP (2001) Superparamagnetic iron oxide contrast 
agents: Physicochemical characteristics and applications in MR imaging. Eur Radiol 
11:2319-2331. 
112. Polyakov VR et al. (2000) Novel tat-peptide chelates for direct transduction of 
Tc-99m and rhenium into human cells for imaging and radiotherapy. Bioconjug Chem 
11:762-771. 
113. Wunderbaldinger P, Josephson L and Weissleder R (2002) Tat peptide directs enhanced 
clearance and hepatic permeability of magnetic nanoparticles. Bioconjug Chem 
13:264-268. 
114. Torchilin VP et al. (1988) Magnetically driven thrombolytic preparation containing 
immobilized streptokinase-targeted transport and action. Haemostasis 18:113-116. 
115. Yoshimoto T et al. (1988) Magnetic urokinase: Targeting of urokinase to fibrin clot. 
Biochem Biophys Res Commun 152:739-743. 
116. Yao F and Eriksson E (2000) Gene therapy in wound repair and regeneration. Wound 
Repair Regen 8:443-451. 
117. Galanis E, Vile R and Russell SJ (2001) Delivery systems intended for in vivo gene 
therapy of cancer: Targeting and replication competent viral vectors. Crit Rev Oncol- 
Hematol 38:177-192. 
118. Buchsbaum DJ and Curiel DT (2001) Gene therapy for the treatment of cancer. Cancer 
Biother Radiopharm 16:275-288. 
119. Murata T et al. (2000) The possibility of gene therapy for the treatment of choroidal 
neovascularization. Ophthalmology 107:1364-1373. 
41 8 Hafeli & Chastellain 
120. Dickson D (1993) UK scientists test liposome gene therapy technique. Nature 365:4. 
121. Mah C et al. (2002) Improved method of recombinant AAV2 delivery for systemic targeted 
gene therapy. Mol Ther 6:106-112. 
122. Scherer F et al. (2002) Magnetofection: Enhancing and targeting gene delivery by magnetic 
force in vitro and in vivo. Gene Ther 9:102-109. 
123. Hughes C, Galea-Lauri J, Farzaneh F and Darling D (2001) Streptavidin paramagnetic 
particles provide a choice of three affinity-based capture and magnetic concentration 
strategies for retroviral vectors. Mol Ther 3:623-630. 
124. Harata K, Matsunaga T and Nagai R (1995) Liposome containing both a magnetosome 
from magnetic bacteria and a gene are useful for studying function and expression of 
genes and in gene therapy. Japan Patent No. 7241192. 
125. Plank C, Anton M, Rudolph C, Rosenecker J and Krotz F (2003) Enhancing and targeting 
nucleic acid delivery by magnetic force. Exp Opin Biol Ther 3:745-758. 
126. Lanza GM et al. (2004) Magnetic resonance molecular imaging with nanoparticles. 
JNucl Cardiol 11:733-743. 
127. Ernst S et al. (2004) Modulation of the slow pathway in the presence of a persistent 
left superior caval vein using the novel magnetic navigation system Niobe. Europace 
6:10-14. 
128. Riffle JS, O'Brien KW, Hafeli UO, Bardenstein D and Dailey JP (2003) Magnetite-based 
polysiloxane fluids for ocular therapies, DARPA Biomagnetics meeting, San Diego, 25. 
129. Riffle JS, Phillips JP and Dailey JP (2002) Magnetic fluids. U.S. Patent No. 6,464,968 B2 
(Oct 15, 2002). 
19 
DQAsomes as Mitochondria-Specific 
Drug and DNA Carriers 
Volkmar Weissig 
1. Introduction 
DQAsomes (i.e. de^ualinium based liposome-like vesicles; pronounced dequasomes) 

have been proposed in 1998 as the first mitochondria-specific colloidal drug and 
DNA delivery system.1 These unique mitochondria-targeted drug carriers have 
been designed based on the intrinsic mitochondriotropism of amphiphilic cations 
with a delocalized charge center, i.e. on cations that accumulate at and inside 
mitochondria of living cells, in response to the mitochondrial membrane potential. 
Prerequisite for creating this system was the distinct self-assembly behavior of 
dicationic quinolinium derivatives, which are mitochondriotropic cations resembling 
"bola"-form electrolytes, i.e. they are symmetrical molecules with two charge 
centers separated by a hydrophobic chain at a relatively large distance. Such "bola"- 
form like amphiphiles form upon sonication of aqueous suspensions cationic vesicles 
("bolasomes") are termed "DQAsomes" when prepared from dequalinium 
salts.1 
The need for mitochondria-specific delivery systems arises from the central role 
mitochondria play in a multitude of metabolic pathways. Mitochondria are vital for 
the cell's energy metabolism and for the regulation of programmed cell death. In 
addition, mitochondria are critically involved in the modulation of intracellular calcium 
concentration and the mitochondrial respiratory chain is the major source of 
damaging reactive oxygen species. Consequently, mitochondrial dysfunction either 
419 
420 Weissig 
causes or at least contributes to a large number of human diseases. Malfunctioning 
mitochondria are found in several adult-onset diseases including diabetes, cardiomyopathy, 
infertility, migraine, blindness, deafness, kidney and liver diseases 
and stroke. The accumulation of somatic mutations in the mitochondrial genome 
has been suggested to be involved in aging, age-related neurodegenerative diseases, 
neuromuscular diseases, as well as in cancer. Consequently, mitochondria 
are increasingly recognized as a prime target for pharmacological intervention.2-5 
The development of methods for selectively delivering biologically active 
molecules to the site of mitochondria, along with the identification of new mitochondrial 
molecular drug targets, will potentially launch new therapeutic approaches 
for the treatment of mitochondria-related diseases, based on either the selective 
protection, repair or eradication of cells.6-9 
2. The Self Assembly Behavior of Bis Quinolinium Derivatives 
2.1. Monte Carlo computer simulations 
Dequalinium salts (Fig. 1A) are dicationic mitochondriotropic compounds resembling 
bola form electrolytes, i.e. they are symmetrical molecules with two charge 
centers separated at a relatively large distance. Such symmetric bola-like structures 
are known from archaeal lipids, which usually consist of two glycerol backbones 
connected by two hydrophobic chains.10 The self-assembly behavior of bipolar 
lipids from Archaea has been extensively studied (reviewed by Gambacorta et al.n). 
It has been shown that these symmetric bipolar archaeal lipids can self-associate 
into mechanically very stable monolayer membranes. 
The most striking structural difference between dequalinium and archaeal 
lipids lies in the number of bridging hydrophobic chains between the polar head 
groups. Contrary to the common arachaeal lipids, in the dequalinium molecule, 
there is only one alkyl chain that connects the two cationic hydrophilic head groups. 
Monte Carlo simulations were applied to a system of bola form amphiphiles in 
a coarse-grained model, in which the amphiphilic molecules consist of connected 
segments with each segment of the chain representing several atoms of a real 
amphiphilic molecule.12 All segments of the coarse grained model are therefore 
either head-like (hydrophilic) or tail-like (hydrophobic) as shown in Fig. IB. 
The formation of molecular aggregates was studied by employing a sequence of 
lattice simulations in an NVT ensemble (constant number of particles, N, constant 
volume, V, constant temperature, T), starting from an isotropic three-dimensional 
distribution of model molecules. The unoccupied lattice sites were considered 
water-like, i.e. hydrophilic. All computer simulations were done at reduced temperatures 
T* = ksT/s and interactions were modeled based on nearest neighbor 
repulsions e (in units of kBT) between hydrophobic tail segments and hydrophilic 
DQAsomes as Mitochondria-Specific Drug and DNA Carriers 421 
- I I - K 
Fig. 1. (A) Chemical structure of dequalinium. (B) Dequalinium after coarse graining.12 
(C) Snapshot from Monte Carlo Computer Simulation: Left, whole vesicle left; Right, cross 
section. 2 (D) Possible conformation of single-chain bola amphiphiles: Left, stretched (bola); 
Right, bended (horse shoe). 
heads. At T* = 0.925 and at 10 voI% amphiphiles (i.e. 10% of all lattice sites were 
occupied with amphiphilic molecules), self-assembled vesicular structures could 
be found, as shown in the snapshot in Fig. IC. Monte Carlo Simulations were also 
used to predict the conformational state of dequalinium within a self-assembled 
structure. While the stretched conformation (Fig. ID, left) would give rise to the 
formation of a monolayer, assuming the horseshoe conformation (Fig. ID, right), 
it would result in the formation of a bilayer. It was found that both conformations 
could theoretically co-exist, although the balance between them appeared to be 
temperature dependent. 
2.2. Physico-chemical characterization 
The self-assembly behavior, as predicted by Monte Carlo Computer Simulation, 
was confirmed by electron microscopy (Fig. 2) and photon correlation spectroscopy 
(Fig. 3).1 It was found that dequalinium forms upon sonication spheric appearing 
aggregates with a diameter between approximately 70 and 700 nm. Freeze fracture 
images (Fig. 2, panel C) show both convex and concave fracture faces. These images 
422 Weissig 
Fig. 2. Electron photomicrograph of DQAsomes. Panel A, negatively stained; Panel B, 
rotary shadowed; Panel C, freeze fractured.1 
30- 
& 20 
10- 
10 100 1000 
Size (nm) 
Fig. 3. Size distribution of DQAsomes.13 
strongly indicate the liposome-like aggregation of dequalinium. Negatively stained 
samples (Fig. 2, panel A) demonstrate that the vesicle is impervious to the stain and 
appears as a clear area surrounded by stain with no substructure visible. Rotary 
shadowed vesicles (Fig. 2, panel B) appeared very electron dense without showing 
any substructure. 
2.3. Structure activity relationship studies 
To define relationships between the structure of dequaMnium-like bola amphiphiles 
and their ability to form bolasomes, the self-assembly behavior of nine dequalinium 
derivatives with varying hydrophilic head groups and different hydrophobic tail 
segments was evaluated.14 It was found that the methyl group in ortho position to 
DQAsomes as Mitochondria-Specific Drug and DNA Carriers 423 
the quaternary nitrogen at the quinolinium ring system seems to play an essential 
role in the self-assembly behavior of these bola amphiphiles; this seems surprising, 
considering the bulky and hydrophobic nature of this group. While the removal of 
this methyl group significantly impairs the stability of formed vesicles, replacing 
the methyl group by an aliphatic ring system (Fig. 4) confers unexpected superior 
vesicle forming properties to this bola amphiphile. Vesicles made from this cyclohexyl 
derivative of dequalinium are contrasted with vesicles made from dequalinium, 
with a very narrow size distribution of 169 ± 50 nm which hardly changes at 
all, even after storage at room temperature for over 5 months. In contrast to vesicles 
made from dequalinium, bolasomes made from the cyclohexyl derivative are 
also stable upon dilution of the original vesicle preparation. While dequaliniumbased 
bolasomes, slowly disintegrate over a period of several hours upon dilution, 
bolasomes made from the cyclohexyl compound do not show any change in size 
distribution following dilution. It appears that bulky aliphatic residues, attached 
to the quinolinium heterocycle, favor self-association of the planar ring system. 
It has therefore been speculated that the bulky group sterically prevents the free 
rotation of the hydrophilic head of the amphiphile around the CH2 - axis (Fig. 5), 
thus contributing to improved intermolecular interactions between the amphiphilic 
monomers.14 
Fig. 4. Structure of the cyclohexyl derivative of dequalinium. 
Fig. 5. Schematic illustration of the stabilizing effect of the cyclohexyl ring system (black 
circles). 
424 Weiss ig 
3. DQAsomes as Mitochondrial Transfection Vector 
The number of diseases found to be associated with defects of the mitochondrial 
genome has grown significantly since 1988. Despite major advances in understanding 
mtDNA, defects at the genetic and biochemical level, there is no satisfactory 
treatment available for a vast majority of patients. Objective limitations of conventional 
biochemical treatment, for patients with defects of mtDNA, warrant 
the exploration of gene therapeutic approaches. Two different strategies for mitochondrial 
gene therapy are imaginable.15 The first involves expressing a wild-type 
copy of the defective gene in the nucleus, with cytoplasmic synthesis and subsequent 
targeting of the gene product to the mitochondria ("allotopic expression"). 
Besides the different codon usage in mitochondria, however, there are possibly 
four major difficulties in adapting this nuclear-cytosolic approach for mitochondrial 
gene therapy to mammalian cells, as recently reviewed by D'Souza.16 Firstly, 
the majority of mtDNA defects involve tRNAs, and to date, no natural mechanism 
has been reported for the mitochondrial uptake of cytosolic tRNAs in mammalian 
cells. Secondly, it is generally agreed that the 13 proteins encoded for by 
mtDNA are very hydrophobic peptides, which would not be readily imported 
by the mitochondrial protein import machinery. However, since the 13 mitochondrial 
coded proteins are not equally hydrophobic, the allotopic expression of at 
least some of the peptides appears as possible.17 Thirdly, it has been hypothesized 
that some of the proteins encoded by the mitochondrion may potentially be 
toxic if synthesized in the cytosol.18 Fourthly, according to a hypothesis termed colocation 
for redox regulation19, the co-location of mtDNA and its products may 
be essential for the rapid control of gene expression by the redox state in the 
mitochondrial matrix. Considering all the problems associated with the nuclearcytosolic 
approach, the development of methods for the direct transfection of 
mitochondria as an alternative approach towards mitochondrial gene therapy is 
warranted. 
Based on the intrinsic mitochondriotropism of dequalinium salts and the ability 
of dequalinium-based vesicles, i.e. DQAsomes, to bind and condense pDNA, 
a strategy for the direct transfection of mitochondria within living mammalian 
cells has been proposed.20'21 This new strategy involves the transport of a DNAmitochondrial 
leader sequence peptide conjugate to mitochondria using cationic 
mitochondriotropic vesicles (DQAsomes), the liberation of this conjugate from the 
cationic vector upon contact with the mitochondrial outer membrane, followed 
by DNA uptake via the mitochondrial protein import machinery. In a series of 
papers,22-27 it was then shown that DQAsomes indeed fulfill all pre-requisites for 
DQAsomes as Mitochondria-Specific Drug and DNA Carriers 425 
a mitochondria-specific DNA delivery system: 
• DQAsomes bind pDNA forming so-called DQAplexes and protect the DNA from 
nuclease digestion. 
• The cytotoxicity of DQAsomes and of DQAsome/pDNA complexes is comparable 
to non-viral transfection vectors, which are already being used in clinical 
trials. 
• DQAsomes mediate the cellular uptake of bound pDNA, most probably via nonspecific 
endocytosis. 
• DQAsomes are endosomolytically active, thereby presumably contributing to an 
early endosomal release of the DQAsome/pDNA complex. 
• DQAplexes do not release pDNA upon contact with anionic phospholipids from 
the inner cytoplasmic membrane. 
• DQAplexes release pDNA upon contact with mitochondria-like membranes, as 
well as upon contact with whole isolated mitochondria. 
• Tested under identical experimental conditions, DQAsomes were shown to transport 
pDNA as well as oligonucleotides to the site of mitochondria, while lipofectin 
was demonstrated to deliver pDNA and oligonucleotides towards the nucleus. 
• Plasmid DNA dissociates from DQAplexes upon contact with mitochondria 
within living mammalian cells. 
Perhaps the most surprising finding among the above listed results is the selective 
DNA release from DQAplexes upon contact with different membranes. Why do 
anionic phospholipids such as phosphatidylserine displace pDNA from lipofectin 
(as shown by Xu and Szoka28), but not from DQAplexes, and why do DQAplexes 
in turn become destabilized upon contact with mitochondrial membranes? When 
looking at data obtained from studies with living mammalian cells,23 it appears 
reasonable to assume that dequalinium molecules could be pulled into the mitochondrial 
matrix in response to the high mitochondrial membrane potential (as 
demonstrated in 1987),29 which in turn might lead to the destabilization of the 
DQAsome/pDNA complex. However, the first detailed study, which demonstrated 
the selective DNA release from DQAplexes, was performed using membranemimicking 
liposomes (Fig. 6). As a model for the intracellular release of DNA 
from DQAsomes, the capacity of anionic liposomes to displace the DNA from 
its cationic carrier was studied. The association of DNA with the cationic carrier 
was assessed by employing SYBR™ Green I. The fluorescence signal of this dye is 
greatly enhanced when bound to DNA. Non-binding results in loss of fluorescence. 
It can be clearly seen that in the vicinity of a 1/1 charge ratio, DQAsomes do not 
release any DNA in the presence of cytoplasmic membrane mimicking liposomes 
426 Weissig 
100 1-1 
.< 80 
1 
« 60 1 
40 1 
- DQAsomes 
(-)/(+) 
20 
\ i Anionic liposomes 
* " i 1 
0 500 1000 
Time [sec] 
Fig. 6. Effect of anionic liposomes on DNA release from DQAsome/pDNA complexes. 
DNA was preincubated with SYBR until stabilization of the signal, followed by adding 
(indicated by arrow) the minimal amount of DQAsomes necessary to decrease the signal 
to background level. Anionic liposomes were then injected (arrow) at an anionic to cationic 
charge ratio (-)/(+) as shown. The displacement of DNA from its carrier is indicated by 
the increase of the fluorescence signal. CPM, cytoplasma membrane like liposomes; IMM, 
inner mitochondrial membrane like liposomes; OMM, outer mitochondrial membrane like 
liposomes.22 
(CPM), not even at a 1.4 fold excess of anionic charge. However, with a similar 
charge excess of anionic liposomes to cationic DQAsomes, 1.6 and 1.7 respectively, 
inner and outer mitochondrial membrane mimicking liposomes (IMM and OMM, 
respectively) are able to displace up to 75% of the DNA from its DQAsomal carrier. 
In agreement with these data, it was found that for the complete liberation of 
DNA from DNA/DQAsome complex, a fourfold excess of dicetylphosphate and 
an eightfold excess of phosphatidylserine, respectively, are necessary. 
The finding that CPM liposomes, at an anionic to cationic charge ratio of 0.82, 
displace up to 75% of the DNA from lipofectin, which was used as a control, do not 
liberate any DNA from DQAsomes even at a slight excess of anionic charge, leads 
to the conclusion that besides the charge ratio, other factors may play an important 
role in the mechanism of DNA release from lipid/DNA complexes. This conclusion 
is being further supported by Xu and Szoka's observation28 that ionic water soluble 
molecules such as ATP, tRNA, DNA, poly(glutamic acid), spermidine and histone 
DQAsomes as Mitochondria-Specific Drug and DNA Carriers 427 
do not displace DNA from the cationic lipid/DNA complex, even at a 100-fold 
charge excess (—/+). In their model for the post-endocytotic release of DNA from 

cationic carriers, they assume the formation of a charge neutral ion pair between 
cationic and anionic lipid, which ultimately results in the displacement of the DNA 
from the cationic lipid and the release of DNA into cytoplasm: 
liposome(+)/DNA(_) + liposome^' <.> [lipid(+)/lipid(_)]liposome + DNA(_) 
According to this equation,22 it seems obvious that an additional gain of free energy 
is obtained by hydrophobic interactions between anionic and cationic lipids, i.e. formation 
of charge neutral liposomes. Considering that there is no difference in the net 
charge between both sides of the equation, the mixed liposome formation should be 
the only driving force leading to DNA release from its lipidic carrier. Intriguingly, 
it was found earlier13 that in physiological solutions, it is not possible to incorporate 
dequalinium into liposomes made of lecithin and lecithin/phosphatidylserine 
respectively. This indicates a very restricted ability of dequalinium to mix with 
phospholipids, which would cause the (assumed) equilibrium in the above equation 
to be on the left side. It was therefore concluded that the miscibility between 
the cationic lipid and the anionic agent used (by nature or by man) to displace the 
DNA is of significant importance.22 
The general feasibility of the DQAsome-based strategy for transfecting mitochondria 
within living mammalian cells, involving pDNA-MLS peptide conjugates, 
has most recently been demonstrated utilizing confocal fluorescence microscopy.30 
It should be noted that the use of physico-chemical methods is, by far, still the only 
way to demonstrate the import of transgene DNA into the mitochondrial matrix 
in living mammalian cells. The complete lack of a mitochondria-specific reporter 
plasmid designed for mitochondrial expression, severely hampers all current efforts 
towards the development of effective mitochondrial expression vectors. While any 
new non-viral transfection system (i.e. cationic lipids, polymers and others) aimed 
at the nuclear-cytosolic expression of proteins can be systematically tested and subsequently 
improved by utilizing any of the many commercially available reporter 
gene systems, such a methodical approach to develop mitochondrial transfection 
systems is currently impossible. A series of papers by Charles Coutelle's laboratory 
describe the principal approach for the design of a mitochondria-specific reporter 
systems.31-33 However, no such system has yet become commercially available. 
It should also be noted that the functional expression of Coutelle's mitochondriaspecific 
expression systems inside the mitochondrial matrix has not been demonstrated 
yet. Thus, evaluating the effectiveness of mitochondria-specific systems 
in delivering DNA into mitochondria depends largely on the physical tracking 
of DNA. 
428 Weissig 
Fig. 7. Confocal fluorescence images of BT20 cells stained with mitotracker (red) after 
exposure for lOhrs to DNA( green) complexed with C-DQAsomes. Left column: circular 
MLSpDNA conjugate, right column: linearized MLS-pDNA conjugate. Top row (A and B): 
red channel, middle row (C and D): green channel, bottom row (E and F): corresponding 
overlaid images. 
Figure 7 shows confocal fluorescence micrographs of cells incubated with MLSpDNA 
conjugates, which were vectorized with vesicles made from the cyclohexyl 
derivative of dequalinium (C-DQAsomes). For the cell exposures imaged in the 
left column (panels A, C and E) the non-restricted, i.e. circular form of pDNA was 
used, while for the experiments pictured in the right column (panels B, D and F), 
the plasmid DNA was linearized before DQAplex formation. The characteristic red 
mitochondrial staining pattern (panels A and B) shows the functional viability of 
the imaged cells and the intracellular green fluorescence (panels C and D) demonstrates 
efficient cell internalization of the fluorescein labeled DNA. The green and 
DQAsomes as Mitochondria-Specific Drug and DNA Carriers 429 
red fluorescence channels were then overlaid to produce the composite image seen 
in panels E and F, where the regions of true co-localization of red and green fluorescence 
were pseudo-colored in white for better visualization. Strikingly, in the 
overlaid images, there is hardly any green fluorescence detectable. Nearly all areas 
of green fluorescence in panels C and D appeared as white areas in panels E and F, 
strongly suggesting that almost the entire DNA has been delivered not only towards 
mitochondria, but also into the organelle. However, whether all or at least a portion 
of the pDNA has actually entered the mitochondrial matrix, i.e. has crossed 
both mitochondrial membranes, and therefore would potentially be accessible to 
the mitochondrial transcription machinery, remains yet to be determined. 
4. DQAsomes as Carriers of Pro-apoptotic Drugs 
Dysregulation of the apoptotic machinery is generally accepted as an almost universal 
component of the transformation process of normal cells into cancer cells 
and a large body of experimental data demonstrates that mitochondria play a key 
role in the complex apoptotic mechanism. Consequently, any therapeutic strategy 
aimed at specifically triggering apoptosis in cancer cells is believed to have 
potential therapeutic effect.34,35 Several clinically approved drugs such as VP-16 
(etoposide), arsenite and vinorelbine, as well as an increasing number of experimental 
anticancer drugs (reviewed by Constantini et al.36), such as betulinic acid, 
lonidamine, ceramide and CD437 have been found to act directly on mitochondria, 
resulting in triggering apoptosis. In order to maximize the therapeutic potential of 
such anticancer drugs, which are known to act at or inside mitochondria, the use of 
DQAsomes as a mitochondria-specific drug delivery system has been proposed.37 
Hypothetically, DQAsome-based anticancer chemotherapy entails features 
which would make it putatively superior to conventional chemotherapeutic 
approaches on the cellular, as well as the subcellular level: 
Firstly, the delivery of drugs known to act directly on mitochondria may trigger 
apoptosis in circumstances in which conventional drugs fail to act, because endogenous, 
"upstream of mitochondria" apoptosis induction pathways are disrupted.36 
Secondly, transporting the cytotoxic drug to its intracellular target could overcome 
multi-drug resistance by hiding the drug inside the delivery system until 
it becomes selectively released at the particular intracellular site of action, i.e. 
mitochondria. 
Thirdly, many carcinoma cells, including human breast adenocarcinomaderived 
cells, have an elevated plasma membrane potential relative to their normal 
parent cell lines in addition to the higher mitochondrial membrane potential.29,39-43 
They could provide the basis for a double-targeting effect of DQAsomes, i.e. on the 
cellular level (normal cells vs. carcinoma cells), and on the sub-cellular level (mitochondria 
versus nucleus). 
430 Weissig 
First data involving the encapsulation of anticancer drags into DQAsomes have 
been published most recently.38 In this study, paclitaxel was chosen as a model compound. 
Paclitaxel is known as a potent antitubulin agent used in the treatment of 
malignancies.44 Its therapeutic potential, however, is limited due to a very narrow 
span between the maximal tolerated dose and intolerable toxic levels. In addition, 
its poor aqueous solubility requires the formulation of emulsions containing 
Cremophor EL®, an oil of considerable toxicity by itself.45 Recently, it has been 
demonstrated that clinically relevant concentrations of paclitaxel target mitochondria 
directly and trigger apoptosis by inducing cytochrome c release in a permeability 
transition pore (PTP)-dependent manner.46 This mechanism of action is known 
from the other pro-apoptotic, directly on mitochondria acting agents.47 A 24-hour 
delay between the treatment with paclitaxel46 or with other PTP inducers,47'48 and 
the release of cytochrome c in cell-free systems, compared with intact cells, has been 
explained by the existence of several drug targets inside the cell, making only a subset 
of the drug available for mitochondria.46 Consequently, paclitaxel was considered 
a prime candidate to benefit from a mitochondria-specific drug delivery system 
such as DQAsomes. It was demonstrated that paclitaxel can be incorporated into 
DQAsomes at a stoichiometric molar ratio of 1 paclitaxel to 2 dequalinium. Considering 
the known spherical character of DQAsomes, the results of an electron 
microscopic (EM) analysis of dequasomal incorporated paclitaxel, however, seem 
rather surprising. The transmission EM image (Fig. 8, left panel) and the cryo-EM 
image (Fig. 9) of an identical sample show a remarkable conformity worm- or rodlike 
structures approximately 400 nm in length, the size of which could also be 
confirmed by the size distribution analysis shown in Fig. 8, right panel. The molecular 
structureof this worm-like complex remains to be determined; nevertheless, 
"Welgfrt »is*rit>u*Io«i Analysis 
9 0 % 
1 0 % 
20O «O0 1O0O 
Size [nm| 
Fig. 8. Left panel: Transmission electron microscopic image (uranyl acetate staining) of 
DQAsomal incorporated paclitaxel (0.67 mol taxol/mol dequalinium); Right panel: Size distribution 
analysis of identical preparation shown in left panel.38 
DQAsomes as Mitochondria-Specific Drug and DNA Carriers 431 
,1 ;- 
* : , . , 
'*>*TW 
Fig. 9. Cryo-electron microscopic image of DQAsomal incorporated paclitaxel (0.67 mol 
dequalinium/mol paclitaxel).38 
the formation of worm-like micelles as described for self-assembling amphiphilic 
block co-polymers49 appears possible. 
In a preliminary study, paclitaxel-loaded DQAsomes were tested for their ability 
to inhibit the growth of human colon cancer cells in nude mice.38 For controls 
with free paclitaxel, the drug was suspended in 100% DMSO at 20 mM, stored at 4°C 
and immediately diluted in warm medium before use. In all controls, the respective 
dose of free paclitaxel and empty DQAsomes was adjusted according to the dose 
of paclitaxel and dequalinium given in the paclitaxel-loaded DQAsome sample. 
Due to the lack of any inhibitory effect on tumor growth, the dose was tripled 
after 1.5 weeks. Figure 10 shows that at concentrations where free paclitaxel and 
I 
! 
i 
t 
1400 
1200 
1000 
800 
600 
400 
200 
0 
» 
•V" 
••''•* -S 
.•// >*"-- x*r 
1 ! 1 
- * - Hepes Buffer 
- • - Free Paclitaxel 
- * - Empty DQAsomes 
- • - Paclitaxel-loaded 
DQAsomes 
10 20 
Day after tumor implantation 
30 
Fig. 10. Tumor growth inhibition study in nude mice implanted with human colon cancer 
cells. The mean tumor volume from each group was blotted against the number of days. Each 
group involved 8 animals. For clarity, error bars were omitted. Note that after 1.5 weeks the 
dose, normalized for paclitaxel, was tripled in all treatment groups.38 
432 Weissig 
empty DQAsomes do not show any impact on tumor growth, paclitaxel-loaded 
DQAsomes (with paclitaxel and dequalinium concentrations identical to controls) 
seem to inhibit the tumor growth by about 50%. Correspondingly, the average 
tumor weight in the treatment group, after sacrificing the animals 26 days, later 
was approximately half of that in all controls. 
Although this result seems to suggest that DQAsomes might indeed be able 
to increase the therapeutic potential of paclitaxel, the preliminary character of this 
first in vivo study has to be emphasized. Experiments to optimize the treatment 
protocol are ongoing in the author's laboratory. 
5. Summary 
Since their initial description in 1998, DQAsomes and DQAsome-like vesicles have 
been established as the first mitochondria-targeted colloidal delivery system, capable 
of transporting plasmid DNA as well as small drug molecules towards mitochondria 
within living mammalian cells. The further exploration of this unique 
mitochondriotropic delivery system will introduce new ways for the treatment of 
cancer and for the therapy of a multitude of mitochondrial diseases. 
Acknowledgments 
I am grateful to Prof. V. P. Torchilin for many helpful discussions and for his strong 
and continuous support of my work. I also would like to sincerely thank my graduate 
students, Gerard D'Souza, Shing-Ming Cheng, Sarathi Boddapati and Eyad 
Katrangi, whose experimental work has made the writing of this chapter possible. 
I am obliged to the Muscular Dystrophy Association (Tucson, AZ), the United 
Mitochondrial Disease Foundation (Pittsburgh, PA), MitoVec, Inc. (Boston, MA) 
and Northeastern University (Boston, MA) for the financial support I received from 
these organizations during the last four years. 
References 
1. Weissig V, Lasch J, Erdos G, Meyer HW, Rowe TC and Hughes J (1998) DQAsomes: 
A novel potential drug and gene delivery system made from Dequalinium. Pharm Res 
15:334-337. 
2. Smith RA, Porteous CM, Gane AM and Murphy MP (2003) Delivery of bioactive 
molecules to mitochondria in vivo. Proc Natl Acad Sci USA 100:5407-5412. 
3. Murphy MP and Smith RA (2000) Drug delivery to mitochondria: The key to mitochondrial 
medicine. Adv Drug Del Rev 41:235-250. 
4. Muratovska A, Lightowlers RN, Taylor RW, Wilce 1A and Murphy MP (2001) Targeting 
large molecules to mitochondria. Adv Drug Del Rev 49:189-198. 
DQAsomes as Mitochondria-Specific Drug and DNA Carriers 433 
5. Szewczyk A and Wojtczak L (2002) Mitochondria as a pharmacological target. Pharmacol 
Rev 54:101-127. 
6. Weissig V (2003) Mitochondrial-targeted drug and DNA delivery. Crit Rev Ther Drug 
Can Syst 20:1-62. 
7. Weissig V, Cheng S-M and D'Souza G (2004) Mitochondrial Pharmaceutics. Mitochondrion 
3:229-244. 
8. Weissig V (2005). Targeted drug delivery to mammalian mitochondria in living cells. 
Exp Opin Drug Del 2:89-102. 
9. Weissig V, Boddapati SV, D'Souza GGM and Cheng SM (2004) Targeting of lowmolecular 
weight drugs to mammalian mitochondria. Drug Des Rev 1:15-28. 
10. De Rosa M, Gambacorta A and Gliozi A (1986) Structure, biosynthesis, and physicochemical 
properties of archaebacterial lipids. Microbiol Rev 50:70-80. 
11. Gambacorta A, Gliozi A and De Rosa M (1995) Archaeal lipids and their biotechnological 
applications. World J Microbiol Biotechnol 11:115-131. 
12. Weissig V, Mogel HJ, Wahab M and Lasch J (1998) Computer simulations of DQAsomes. 
Proc Intl Symp Control Rel Bioact Mater 25:312. 
13. Weissig V, Lizano C and Torchilin VP (1998) A micellar delivery system for 
dequalinium — A lipophilic cationic drug with anticarcinoma activity. / Lipos Res 
8:391-400. 
14. Weissig V, Lizano C, Ganellin CR and Torchilin VP (2001) DNA binding cationic bolasomes 
with delocalized charge center: A structure-activity relationship study. STP 
Pharma Sci 11:91-96. 
15. Chrzanowska-Lightowlers ZM, Lightowlers RN and Turnbull DM (1995) Gene therapy 
for mitochondrial DNA defects: Is it possible? Gene Ther 2:311-316. 
16. D'Souza GGM and Weissig V (2004) Approaches to mitochondrial gene therapy. Curr 
Gene Ther 4:317-328. 
17. Manfredi G, Fu J, Ojaimi J, Sadlock JE, Kwong JQ, Guy J and Schon EA (2002) Rescue 
of a deficiency in ATP synthesis by transfer of MTATP6, a mitochondrial DNA-encoded 
gene, to the nucleus. Nat Genet 30:394-399. 
18. Jacobs HT (1991) Structural similarities between a mitochondrially encoded polypeptide 
and a family of prokaryotic respiratory toxins involved in plasmid maintenance suggest 
a novel mechanism for the evolutionary maintenance of mitochondrial DNA. /Mo/ Evol 
32:333-339. 
19. Allen JF (2003) The function of genomes in bioenergetic organelles. Philos Trans R Soc 
Lond B Biol Sci 358:19-37. 
20. Weissig V and Torchilin VP (2000) Mitochondriotropic cationic vesicles: A strategy 
towards mitochondrial gene therapy. Curr Pharm Biotechnol 1:325-346. 
21. Weissig V and Torchilin VP (2001) Towards mitochondrial gene therapy: DQAsomes as 
a strategy. / Drug Targ 9:1-13. 
22. Weissig V, Lizano C and Torchilin VP (2000) Selective DNA release from DQAsome / DNA 
complexes at mitochondria-like membranes. Drug Del 7:1-5. 
23. D'Souza GG, Rammohan R, Cheng SM, Torchilin VP and Weissig V (2003) DQAsomemediated 
delivery of plasmid DNA toward mitochondria in living cells. / Control Rel 
92:189-197. 
434 Weissig 
24. Lasch J, Meye A, Taubert H, Koelsch R, Mansa-ard J and Weissig V (1999) Dequalinium 
vesicles form stable complexes with plasmid DNA which are protected from DNase 
attack. Biol Chem 380:647-652. 
25. Weissig V, D'Souza GG and Torchilin VP (2001) DQAsome/DNA complexes release 
DNA upon contact with isolated mouse liver mitochondria. / Control Rel 75:401-408. 
26. Weissig V, Seibel P, Seibel M and Torchilin V P (2001) Binding and release of DNApeptide 
conjugates by cationic mitochondriotropic vesicles (DQAsomes). Proc Intl 
Symp Control Rel Bioact Mater 28:850. 
27. D'Souza GG, Boddaparti S and Weissig V (2004) Proc Intl Symp Control Rel Bioact Mater 
31. 
28. Xu Y and Szoka FC, Jr. (1996) Mechanism of DNA release from cationic liposome/DNA 
complexes used in cell transfection. Biochemistry 35:5616-5623. 
29. Weiss MJ, Wong JR, Ha CS, Bleday R, Salem RR, Steele GD, Jr. and Chen LB (1987) 
Dequalinium, a topical antimicrobial agent, displays anticarcinoma activity based on 
selective mitochondrial accumulation. Proc Natl Acad Sci USA 84:5444-5448. 
30. D'Souza GG, Boddapati S, Lightowlers RN and Weissig V (2005) Mitochondriotropic 
vesicles deliver mitochondrial leader peptide conjugates of circular and linear doublestranded 
DNA into mammalian mitochondria. Proc Intl Symp Control Rel Bioact Mater 
32: 
31. Wheeler VC, Prodromou C, Pearl LH, Williamson R and Coutelle C (1996) Synthesis of 
a modified gene encoding human ornithine transcarbamylase for expression in mammalian 
mitochondrial and universal translation systems: A novel approach towards 
correction of a genetic defect. Gene 169:251-255. 
32. Wheeler VC, Aitken M and Coutelle C (1997) Modification of the mouse mitochondrial 
genome by insertion of an exogenous gene. Gene 198:203-209. 
33. Bigger BW, Tolmachov O, Collombet JM, Fragkos M, Palaszewski I and Coutelle C (2001) 
An araC-controlled bacterial ere expression system to produce DNA minicircle vectors 
for nuclear and mitochondrial gene therapy. / Biol Chem 276:23018-23027 
34. Ferreira CG, Epping M, Kruyt AE and Giaccone G (2002) Apoptosis: Target of cancer 
therapy. Clin Cancer Res 8:2024-2034. 
35. Reed JC (1999) Dysregulation of apoptosis in cancer. / Clin Oncol 17:2941-2953. 
36. Costantini P, Jacotot E, Decaudin D and Kroemer G (2000) Mitochondrion as a novel 
target of anticancer chemotherapy. / Nat Cancer Inst 92:1042-1053. 
37. Weissig V, Cheng S-M, Pabba S, D'Souza G, Torchilin VP, Schubert R and Kimpfler A 
(2003) A novel strategy for mitochondria-specific delivery of apoptosis-inducing agents: 
DQAsomal incorporated paclitaxel. Proc Intl Symp Control Rel Bioact Mater 30:505. 
38. Cheng SM, Pabba S, Torchilin VP, Fowle W, Kimpfler A, Schubert R and Weissig V 
(2005) Towards mitochondria-specific delivery of apoptosis-inducing agents: DQAsomal 
incorporated paclitaxel. / Drug Del Sci Technol 14. 
39. Modica-Napolitano JS and Aprille JR (1987) Basis for the selective cytotoxicity of rhodamine 
123. Cancer Res 47:4361-4365. 
40. Modica-Napolitano JS and Aprille JR (2001) Delocalized lipophilic cations selectively 
target the mitochondria of carcinoma cells. Adv Drug Del Rev 49:63-70. 
DQAsomes as Mitochondria-Specific Drug and DNA Carriers 435 
41. Modica-Napolitano JS, Koya K, Weisberg E, Brunelli BT, Li Y and Chen LB (1996) 
Selective damage to carcinoma mitochondria by the rhodacyanine MKT-077. Cancer Res 
56:544-550. 
42. Manetta A, Emma D, Gamboa G, Liao S, Berman M and DiSaia P (1993) Failure to 
enhance the in vivo killing of human ovarian carcinoma by sequential treatment with 
dequalinium chloride and tumor necrosis factor. Gynecol Oncol 50:38-44. 
43. Christman JE, Miller DS, Coward P, Smith LH and Teng NN (1990) Study of the selective 
cytotoxic properties of cationic, lipophilic mitochondrial-specific compounds in gynecologic 
malignancies. Gynecol Oncol 39:72-79. 
44. Eisenhauer EA and Vermorken JB (1998) The taxoids. Comparative clinical pharmacology 
and therapeutic potential. Drugs 55:5-30. 
45. Seligson AL, Terry RC, Bressi JC, Douglass JG, 3rd and Sovak M (2001) A new prodrug 
of paclitaxel: Synthesis of Protaxel. Anticancer Drugs 12:305-313. 
46. Andre N, Carre M, Brasseur G, Pourroy B, Kovacic H, Briand C and Braguer D (2002) 
Paclitaxel targets mitochondria upstream of caspase activation in intact human neuroblastoma 
cells. FEBS Lett 532:256-260. 
47. Fulda S, Susin SA, Kroemer G and Debatin KM (1998) Molecular ordering of apoptosis 
induced by anticancer drugs in neuroblastoma cells. Cancer Res 58:4453-4460. 
48. Waterhouse NJ, Ricci JE and Green DR (2002) And all of a sudden it's over: Mitochondrial 
outer-membrane permeabilization in apoptosis. Biochimie 84:113-121. 
49. Discher ED and Eisenberg A (2002) Polymer vesicles. Science 297:967-973. 
This page is intentionally left blank
20 
Liposomal Drug Carriers 
in Cancer Therapy 
Alberto A. Gabizon 
1. Introduction 
In the last two decades, we have witnessed the development of implantable and 
injectable drug delivery systems for applications in the treatment of cancer and 
other diseases. These systems have arisen from various needs: 
1. To provide depot forms of drug administration and more convenient dosing 
schedules. Examples are implantable biodegradable rods for slow release of 
peptides such as LH-RH partial agonists (e.g. goserelin depot), for blockade of 
gonadal production of androgens or estrogens.1 This is a simple and pharmacologically 
effective approach developed for an implantable drug delivery system. 
2. To provide for convenient vehicles of administration for poorly soluble drugs. 
These systems may or may not confer an advantage to the therapeutic index 
of the drug, but their basic "raison d'etre" is to provide a vehicle of injection. 
An example is paclitaxel entrapped in polymerized albumin nanoparticles2 for 
i.v. administration of paclitaxel in cremophor-free form. These should be distinguished 
from simple excipients used as solubilizers (e.g. cremophor in the case 
of paclitaxel), because, in the former, the drug and vehicle are physically in one 
single complex at least during the initial phase in circulation. 
3. To improve the efficacy and reduce the side effects of new and old anticancer 
drugs. Examples include formulations of anthracyclines encapsulated 
in liposomes (e.g. Doxil, Myocet, Daunoxome) or conjugated to polymers.3 
437 
438 Gabizon 
The objective here is to change the pharmacokinetics, biodistribution, and the 
bioavailability profile so as to achieve a positive impact on the drug pharmacodynamics. 
This is the most refined approach to drug delivery and includes 
intravenous administration of a d r u g stably associated to a carrier, with or without 
specificity to a target cancer cell molecule. 
In this review chapter, we will focus on injectable particulate drug delivery 
systems of anticancer drugs (Fig. 1), particularly on liposomes, the most widely 
used drug nanocarrier in cancer. The physico-chemical properties of liposomes are 
discussed elsewhere in this book. Briefly, liposomes are vesicles with an aqueous 
interior surrounded by one or more concentric bilayers of phospholipids with a 
diameter ranging from a minimal diameter of ~30 nm to several microns. However, 
for injectable clinical applications, practically all liposome formulations are in 
the submicron ultrafilterable range (<200 nm size) and can be considered as nanosize 
particulate systems. Liposomes are formed spontaneously when amphiphilic 
lipids such as phospholipids are dispersed in water. The ensuing structures are 
physically stable supramolecular assemblies, and unlike polymerized particles, 
Fig. 1. Cancer therapy and drug delivery systems — Schematic drawing illustrating various 
approaches to smart cancer drug delivery. (1) Targeting of drugs conjugated to antibodies 
or ligands directed to tumor cell-specific surface receptors. (2) Controlled release of drugs 
entrapped in microspheres or nanospheres by diffusion and/or degradation of the particle 
matrix in extracellular fluid. (3) Release of drugs entrapped in phospholipids vesicles 
(liposomes) by leakage and / o r endocytosis and liposome breakdown. (4) Delivery of drugs 
conjugated to polymers by endocytosis and intracellular drug release. These approaches may 
also be combined, for instance, in the case of liposomes targeted with ligands or antibodies 
to tumor cells. [Note: relative scales are disproportional.] 
Liposomal Drug Carriers in Cancer Therapy 439 
they are not covalently bound. Although liposome formation is actually a spontaneous 
process, the current trend is to classify them into a class of pharmaceutical 
devices in the nanoscale range engineered by physical and/or chemical means, and 
referred to as nanomedicines.4 Nanomedicines are a direct result of the application 
of nanotechnology to medicine, and encompass in their wide context, molecular and 
supramolecular devices such as liposomes and other nanoparticulate carriers. In 
fact, liposomes are the first generation of nanosize drug delivery devices approved 
for the treatment of cancer (i.e. Doxil containing doxorubicin) and fungal infections 
(i.e. Ambisome containing amphotericin B). Current liposome formulations represent 
a basic form of nanomedicine involving a slow drug release system, and often 
a passive targeting process known as enhanced permeability and retention (EPR) 
that will be discussed later in this chapter. The field of nanomedicines is rapidly 
evolving and aims at increased sophistication of nanosize devices interacting with 
cellular targets at the nanoscale level. 
2. The Challenge of Cancer Therapy 
Our understanding of the molecular processes underlying the pathologic behavior 
of cancer cells has progressed enormously in the last decade.5 Of particular relevance 
to cancer targeting is the fact that a number of receptors, mostly growth 
factor receptors, have been found to be overexpressed in tumor cells, and to play 
an important role as catalysts of growth. Receptor profiling of tumors may offer 
a potential Achilles heel for targeting specific ligands or antibodies, with or without 
delivery of a cytotoxic drug cargo.6 In addition, the pathophysiology of tumor 
neovasculature and the interaction of tumor with stroma have been recognized as 
processes that play a major role in tumor development. Cancer is ultimately a disease 
caused by somatic gene mutations that result in the transformation of a normal 
cell into a malignant tumor cell. Eventually, the tumor cell phenotype progresses 
along three major steps7: 
1. Increased proliferation rate and/or decreased apoptosis, causing an increase of 
tumor cell mass. 
2. Invasion of surrounding tissues and switch on of angiogenesis. This is a critical 
step that differentiates in situ, non-invasive, tumors with no metastatic potential 
from invasive tumors with metastatic and life-threatening potential. Although 
there is considerable variability, tumors with angiogenic potential become vascularized 
when the cell load reaches an order of 107 cells, equivalent to a nodule 
of ~2 mm diameter. 
3. Metastases, i.e. abnormal migration of tumor cells from the primary tumor site 
via blood vessels or lymphatics to distant organs, with formation of secondary 
tumors. This is most commonly the process that causes death of the host due 
440 Gabizon 
to disruption of the function of vital organs or systems (i.e. brain, lung, liver, 
kidney, bone marrow, coagulation, intestinal passage, and others). 
Despite formidable advances in clinical imaging, the diagnosis of a tumor mass1 
usually requires the presence of a nodule of ~10 mm diameter, representing a cluster 
of 109 cells.2 Since the lethal tumor burden is in the order of 1012 cells in most 
cancer patients, this implies that tumors have already gone through 75% of their 
doubling cell expansion process by the time of clinical diagnosis. As a result, significant 
heterogeneity and phenotypic diversity are already present in most diagnosed 
cancers, posing a major therapeutic challenge due to the development of metastatic 
ability and drug resistance. 
Nowadays, drug-based therapy of cancer is applied in three possible settings: 
• Primary treatment, which is also known as neo-adjuvant or pre-operative treatment. 
In this setting, anti tumor drugs are given prior to potentially curative 
local therapeutic modalities such as surgery or radiotherapy. These patients have 
a primary tumor, but no clinical evidence of distant metastases. Concomitant 
treatments of chemotherapy, or hormonal therapy, with radiotherapy can also 
be included in this category. The goal is to reduce tumor bulk, the risk of tumor 
seeding, and to facilitate surgery or radiotherapy of the primary tumor. 
• Adjuvant treatment. The aim is to eradicate clinically undetectable residual tumor 
cells, presumably left over after surgical removal of the primary tumor. Adjuvant 
treatment is generally applied in patients with a high risk of micrometastases. 
To some extent, the adjuvant approach likens to a black box because all patients 
at high risk are treated without knowing for sure who are the patients harboring 
metastases and who are not. Also, we have no immediate way of knowing 
whether the treatment is effective or not. Only long follow-up periods will reveal 
if cancer will recur in a specific patient. Therefore, the proof of efficacy of adjuvant 
treatment is exclusively statistical. Despite these limitations, it has been demonstrated 
statistically that adjuvant treatment can cure subclinical, micrometastatic 
disease in a fraction of patients with breast cancer, colon cancer, and other tumors, 
who would not be amenable to cure if the disease is to become macroscopic and 
clinically detectable prior to treatment. The evaluation of adjuvant treatment 
effects is complicated by the poorly understood phenomenon of tumor dormancy 
in which tumor cells appear to remain as tiny, quiescent avascular clusters for 
1 This does not apply to superficial skin tumors which can be recognized sometimes when tumors contain 
cell clusters of 107 reaching ~2-3mm diameter. 
Occasionally modern imaging techniques (high resolution CT scan, MRI) can detect smaller (3-5 mm) 
findings with suspected cancer features in asymptomatic individuals. However, it is reasonable to 
assume that non-imaging techniques, for ex. proteomics based, will be needed to safely break through 
the 108-9 cancer cell mass diagnostic threshold. 
Liposomal Drug Carriers in Cancer Therapy 441 
long periods of time. Small, microscopic, tumor cell clusters may get their nutrients 
by diffusion from pre-existing adjacent vessels of normal tissues. Therefore, 
adjuvant therapies specifically directed to tumor vasculature are unlikely to be 
effective against some micrometastases during the avascular phase. 
• Treatment of metastatic disease or neoplastic conditions not amenable to surgical 
or radiotherapeutic eradication. In these cases, chemotherapy is potentially curative 
only in hematological and lymphoid neoplasms, and in a few cases of solid 
tumors such as testicular cancer and choriocarcinoma. In most instances, including 
the most common types of cancer namely breast, prostate, lung, and colon, 
chemotherapy is palliative, i.e. temporary tumor regression and prolongation of 
survival can be achieved, but cure is exceptional and most tumors ultimately 
recur and are lethal. 
Let us now examine the currently available cancer drug armamentarium. Drug 
therapies of cancer can be divided into three major groups: 
1. Cytotoxic agents. As the name implies, these agents are toxic to cells and lack 
tumor cell specificity. They can be divided into three major groups: 
• Agents that damage the DNA template directly or indirectly. 
• Agents that damage the microtubule-based spindle apparatus. 
• Agents that inhibit DNA synthesis (antimetabolites). 
Upon structural damage or arrest of the cell cycle, tumor cells undergo apoptosis 
which is the main form of cell death. Treatment with cytotoxic agents is usually 
referred to as cancer chemotherapy. The use of cytotoxic agents remains the mainstay 
of cancer therapy. It is this group of agents that urgently requires a delivery 
system to improve its tumor specificity, and/or reduce its damage to normal tissues. 
In addition to the lack of specificity of chemotherapeutic (cytotoxic) agents, a 
number of physiologic factors can seriously limit the efficiency of drug distribution 
from plasma to tumors and neutralize their effects. These include competition for 
drug uptake of well-perfused tissues such as liver and kidneys, rapid glomerular 
filtration and urinary excretion of low molecular weight drugs, protein binding 
with drug inactivation (e.g. cisplatin), and stability problems in biological fluids 
(e.g. hydrolysis of nitrosoureas, opening of lactone ring of camptothecin analogs). 
2. Hormonal agents. They are used mainly against breast and prostate cancers. These 
tumors often require estrogen or androgen receptor activation for growth stimulation. 
The hormonal therapies currently in use are mostly based on synthetic 
compounds modeled to block the gonadal or peripheral production of estrogens 
and androgens (i.e. LH-RH partial agonists, aromatase inhibitors) or to compete 
for the tumor cell receptors of these hormones (i.e. anti-estrogens, anti-androgens). 
442 Gabizon 
Corticosteroids and somatostatin analogs can also be included in the category of 
hormonal agents. 
3. Non-cytotoxic agents modifying biological response. These can be classified in at least 
three distinct groups: 
• Antibodies blocking growth factors, growth factor receptors and other cellmembrane 
receptors of tumor cells or supporting stroma (e.g. bevacizumab, 
trastuzumab, cetuximab, rituximab). 
• Agents blocking signal transduction kinases (e.g. gefitinib, imatinib). 
• Cytokines with miscellaneous activities (e.g. interferon-a, interleukin-2). 
3. The Rationale for the Use of Liposomal 
Drug Carriers in Cancer 
The rationale for the use of liposomes in cancer drug delivery is based on the 
following pharmacological principles,8 which are also applicable to non-liposomal 
nanoparticulate drug carriers: 
1. Slow drug release. Drug bioavailability depends on drug release from liposomes. 
Entrapment of drug in liposomes will slow down drug release and reduce renal 
clearance to a variable extent. Slow release may range from a mere blunting of the 
peak plasma levels of free drug, to a sustained release of drug mimicking continuous 
infusion. These pharmacokinetic changes may have important pharmacodynamic 
consequences with regard to toxicity and efficacy of the liposome delivered agents. 
2. Site avoidance of specific tissues. The biodistribution pattern of liposomes may lead 
to a relative reduction of drug concentration in tissues specifically sensitive to the 
delivered drug. This may have implications with regard to the therapeutic window 
of various cytotoxic drugs, such as the cardiotoxic anthracyclines, provided that 
anti tumor efficacy is not negatively affected. 
3. Accumulation in tumors. Prolongation of the circulation time of liposomes results 
in significant accumulation in tissues with increased vascular permeability This 
is often the case of tumors,9 especially in those areas with active neoangiogenesis. 
Tumor localization of long-circulating liposomes, such as pegylated liposomes, 
sometimes referred to as Stealth or sterically-stabilized,10 is a passive targeting effect 
that enables substantial accumulation of liposome-encapsulated drug in the interstitial 
fluid at the tumor site,11 a phenomenon sometimes referred to as enhanced 
permeability and retention (E.P.R.) effect (Fig. 2). 
There are a number of differential effects of physiologic factors on clearance and 
biodistribution of low molecular weight drugs and nanoparticles (see also Table 1): 
• Protein binding. Low molecular weight drugs may be inactivated and/or irreversibly 
bound by plasma proteins, thus reducing the bioavailability towards 
Liposomal Drug Carriers in Cancer Therapy 443 
Fig. 2. Extravasation and release of liposomal drug cargo in tumor interstitial fluid compartment 
— Schematic drawing illustrating the concept of passive targeting of liposomes 
to tumors exploiting the EPR effect. The dots represent the drug molecules encapsulated in 
the liposome water phase. The various steps implied in the targeting process are numerically 
designated from 1 to 5. (1) Liposomes with long-circulating properties are required to 
increase the number of passages through the tumor microvasculature. (2) Increased vascular 
permeability in tumor tissue enables properly downsized liposomes to extravasate and 
reach the tumor interstitial fluid. (3) Because of their limited diffusion capacity, liposomes 
remain in close vicinity to blood vessels. (4) Drug is gradually released from liposomes accumulating 
in the interstitial fluid moving swiftly through the tumor cell layers and entering 
tumor cells. (5) The cytotoxic effect leading to tumor cell death is expected to follow the same 
mechanism known for free drug. \Nole: relative scales are disproportional.] 
cellular target molecules. Cisplatin, a widely used anticancer cytotoxic drug, is 
one such example. In the case of nanoparticles, plasma proteins can adsorb to 
their surface a process known as opsonization that results in tagging the particle 
for recognition and removing it by macrophages. In addition, protein binding to 
the liposome surface may de-stabilize the bilayer and accelerate the leakage of 
liposome contents. PEG coating (pegylation) of liposomes reduces opsonization 
and the effects associated with it. 
• Reticulo-endothelial system (RES) clearance. It is unimportant for low molecular 
weight drugs, but plays a major role in the clearance of nanoparticles reducing the 
fraction available for distribution to tumor tissue. Kupffer cell macrophages lining 
the liver sinusoids remove opsonized liposomes and other nanoparticles from 
circulation, and represent a major factor in the clearance of particulate carriers. 
444 Gabizon 
Table 1 Differential effects of physiologic factors on clearance and biodistribution of low 
molecular weight (MW) drugs and nanoparticles. 
Factor Extravascular Microvascular Glomerular Protein binding R.E.S. 
transport permeability filtration clearance 
(fenestrations) 
Low MW drugs Diffusion Not Filterable Binding and Unimportant 
important inactivation 
Nanoparticles Convection Critical for Non- Opsonization Major 
tissue filterable and clearance 
targeting de-stabilization pathway 
• Glomerular filtration. Unless they become protein-bound, low molecular weight 
drugs can be filtered out by kidney glomeruli. In contrast, liposomes and other 
nanoparticles are non-filterable, because their diameter exceeds the glomerular 
filterable threshold size. 
• Microvascular permeability. Enhanced microvascular permeability with fenestrations 
in capillaries and post-capillary venules is critical for the extravasation of 
nanoparticles from the blood stream to the interstitial fluid of the target tissues. 
The presence of fenestrations is irrelevant for tissue delivery of small molecules. 
• Extravascular transport. Diffusion is the predominant mechanism of transport 
for small molecules. In contrast, convective transport plays a major role in the 
extravascular movement of nanoparticles, for which diffusion rates are very 
slow.12 Large tumors tend to develop high interstitial pressure that reduces the 
rate of convective transport significantly. In fact, in an animal model, it has been 
shown that liposomes accumulate significantly less in larger tumors on a per gram 
tissue basis.13 In agreement with this, large tumor size predicts poor response to 
liposome-delivered chemotherapy in ovarian cancer.14 Table 2 lists a number of 
tumor and liposome factors that play important roles in the delivery of liposomal 
drugs. On the tumor side, a rich blood flow and a highly permeable microvascular 
Table 2 Parameters affecting delivery of liposomal drugs 
to tumors. 
Tumor factors Liposome factors 
• Blood flow • Long circulation time 
• Vascular permeability • Stability (drug retention) 
• Interstitial pressure • Small vesicle size 
• Phagocytic activity • Saturation of the RES 
Liposomal Drug Carriers in Cancer Therapy 445 
bed will increase the probability of liposome deposition, while a high interstitial 
fluid pressure is likely to reduce the movement of molecules and particles 
into the tumor compartment. On the liposome side, avoiding drug leakage and 
prolonging the circulation time will result in more liposomes reaching the tumor 
vascular bed with an intact drug payload, and a small vesicle size will facilitate 
extravasation through the endothelium gaps or fenestrations. There are also data 
indicating that saturation of the RES will prolong circulation time and indirectly 
enhance liposome deposition in tumors.15 
4. Liposome Formulation and Pharmacokinetics — 
Stealth Liposomes 
In 1971, Gregoriadis et a\}b published the first research work in which liposomes 
were used as drug carriers for medical applications. This initial study led to growing 
interest in liposomes, and many laboratories began examining liposome pharmacokinetics 
and biodistribution in animals, as well as in vitro stability in serum. The 
early liposome work was mostly based on formulations composed of neutral egg 
lecithin (PC), often in combination with negatively or positively charged lipids. 
These liposomes were found to release rapidly a large fraction of their encapsulated 
contents in circulation. Furthermore, they were quickly removed from the 
circulation by macrophages of the RES. Reformulation with high phase-transition 
temperature (Tm) lipids (distearoyl-PC, dipalmitoyl-PC, sphingomyelin) and addition 
of cholesterol led to improved retention of liposome contents and prolongation 
of circulation time, especially when the vesicles were properly downsized 
to <100nm diameter. However, these relatively improved liposome formulations 
would still accumulate largely in the RES and a greater improvement in circulation 
half-life appeared to be required for cancer targeting. Surface modifications 
of liposomes that could reduce the RES affinity were investigated based on the 
erythyrocyte paradigm, whereby a layer of carbohydrate groups prolongs circulation 
for nearly 3 months. A number of glycolipids such as monosialoganglioside 
(GM1), phosphatidyl-inositol, and cerebroside sulfate, were included in the 
formulations and extended liposome circulation time.17,18'19 However, a major 
advance took place when the hydrophilic polymer polyethylene-glycol (PEG), 
which was known to reduce immunogenicity and prolong circulation time when 
attached to enzymes and growth factors, was introduced into liposomes in the 
early 90s. PEG, which is inexpensive due to easy synthesis and could be prepared 
in high purity and large quantities, had distinct advantages over the other glycolipid 
surface modifiers. Addition of a conjugate of PEG with a lipid anchor, 
distearoyl-phosphatidylethanolamine (PEG-DSPE) to the liposomal formulation 
was shown to prolong liposome circulation time significantly,20,21,22 and formed 
446 Gabizon 
a pivotal element of the pharmaceutical development of the Doxil formulation 
described thereafter in this chapter. Due to their ability to avoid RES clearance 
mechanisms, PEG-coated liposomes have been coined "Stealth"3 liposomes.23 In 
parallel to the development of stable formulations with longer circulation halflives, 
it was soon realized that a prolonged residence time in circulation was a 
critical pharmacokinetic factor for liposome deposition in tumors and that there 
was a strong correlation between liposome circulation time and tumor uptake.18 
A number of studies have addressed the mechanism of liposome accumulation 
in tumors. Microscopic observations with colloidal gold-labeled liposomeS24 and 
morphologic studies with fluorescent liposomes in the skin-fold chamber model25 
have demonstrated that liposomes extravasate into the tumor extracellular fluid 
through gaps in tumor microvessels and are found predominantly in the perivascular 
area with minimal uptake by tumor cells. Studies with ascitic tumors26'27 
demonstrate a steady extravasation process of long circulating liposomes into the 
ascitic fluid, with gradual release of drug followed by drug diffusion into the ascitic 
cellular compartment. The process underlying the preferential tumor accumulation 
of liposomes, as well as other macromolecular and particulate carriers, is known 
as EPR (enhanced permeability and retention) effect.28 This is a passive and nonspecific 
process resulting from increased microvascular permeability and defective 
lymphatic drainage in tumors creating an in situ depot of liposomes in the tumor 
interstitial fluid. Circulating liposomes cross the leaky tumor vasculature, moving 
from plasma into the interstitial fluid of tumor tissue, following convective transport 
and diffusion processes. Although convective transport of plasma fluid also 
occurs in normal tissues, the continuous, non-fenestrated endothelium and basement 
membrane prevents the extravasation of liposomes. EPR is a relatively slow 
process, in which long-circulating liposomes possess a distinct advantage because 
of the repeated passage through the tumor microvascular bed and their high concentration 
in plasma during an extended period of time. 
For any intra-vascular drug carrier device to access the tumor cell compartment 
and interact with tumor cell receptors, it must first cross the vascular endothelium 
and diffuse into the interstitial fluid, since with few exceptions, tumor cells and their 
surface receptors are not directly exposed to the blood stream. Therefore, the EPR 
effect is not only important for the tumor accumulation of non-targeted liposomes, 
but it is also for that of ligand-targeted liposomes. This has led us to postulate 
that the extravasation process is the rate-limiting step of liposome accumulation in 
tumors.29 Experimental data with targeted and non-targeted liposomes have so far 
lent consistent support to this hypothesis.30 
3Stealth is a registered trademark of Alza Corp., Mountain View, CA. 
Liposomal Drug Carriers in Cancer Therapy 447 
In most instances, delivery of drug to tumor cells depends on the release of 
drug from liposomes in the interstitial fluid, since liposomes are seldom taken up 
by tumor cells, unless they are tagged with specific ligands. The factors controlling 
this process and its kinetics are not well understood and may vary among 
tissues, depending on the liposome formulation in question. In the case of remoteloaded 
formulations, e.g. anthracyclines, a gradual loss of the liposome gradient 
retaining the drug, in addition to the disruption of the integrity of the liposome 
bilayer by phospholipases, may be involved in the release process. Uptake by 
tumor-infiltrating macrophages could also contribute to liposomal drug release. 
In any case, once the drug is released from liposomes, it will diffuse freely through 
the interstitial tumor space and reach deep layers of tumor cells. This is an inherent 
advantage of this delivery system as opposed to covalently bound drug-carrier 
systems. It is also a critical factor for the success of the liposomal drug approach, 
since most of the liposomes appear to remain in interstitial spaces immediately 
surrounding the blood vessels,25 and therefore would not be able to interact with 
more than one layer of tumor cells. 
The EPR effect has been confirmed in a variety of implanted tumor models. 
Its validity regarding human tumors, and particularly, cancer metastases, is as yet 
unclear. One concern is that interstitial fluid pressure increases in most tumors 
once they grow beyond a certain size threshold,31 thereby hindering extravascular 
transport and liposome delivery. Unfortunately, there is a paucity of imaging studies 
in cancer patients with radiolabeled liposomes. One of the few studies with 
radiolabeled pegylated liposomes demonstrated significant liposome accumulation 
based on tumor imaging findings in 15 out of 17 patients tested.32 In another 
study, in which tumor metastases and normal muscle tissues of 2 breast cancer 
patients were examined for doxorubicin concentration after injection of pegylated 
liposomal doxorubicin (PLD), liposomal drug was found at 10-fold greater concentration 
in tumor, as compared with muscle.33 Another important piece of work 
in this area is the study of Northfelt et a/.,34 that pointed to an enhanced deposition 
of drug in Kaposi's sarcoma skin lesions of patients receiving PLD, compared 
with the normal skin of the same patients and to doxorubicin concentration in 
Kaposi's sarcoma biopsies of patients receiving free doxorubicin. More imaging 
and drug-carrier biodistribution studies are needed to determine how important 
and frequent is the observation of human tumors with selectively enhanced uptake 
of liposomes. These studies would also enable to determine whether there is a correlation 
between liposome accumulation in tumors and anti tumor response, and 
a need to select patients for liposome-delivered drug therapy based on positive 
liposome tumor imaging. 
448 Gabizon 
5. Preclinical Observations with Liposomal Drug Carriers 
in Tumor Models 
The drug most frequently tested in liposomal formulations is doxorubicin and 
related anthracyclines. The choice of doxorubicin by many of the early research 
groups examining the role of liposomes as drug carriers in cancer chemotherapy, 
stems from its broad spectrum of anti tumor activity on the one hand, and its disturbing 
cumulative dose-limiting cardiac toxicity on the other hand. 
Anthracyclines such as doxorubicin and daunorubicin cause acute toxic side 
effects including bone marrow depression, alopecia, and stomatitis, and are dose 
limited by a serious and mostly irreversible characteristic cardiomyopathy.35 The 
first study describing the encapsulation of anthracyclines into liposomes appeared 
in 1979.36 Work from various research groups followed, supporting the general principle 
that liposomal formulations reduced the toxicity of anthracyclines in animal 
models. 
Using the Stealth technology and an elegant loading mechanism based on 
an ammonium sulfate gradient, a formulation of pegylated liposomal doxorubicin 
(PLD) known as Doxil in the USA (Caelyx in Europe) has been developed. 
The loading mechanism, coined "remote (active) loading", leads to highly efficient 
accumulation of doxorubicin inside the aqueous phase (~ 15,000 doxorubicin 
molecules/vesicle), where the drug forms a crystalline-like precipitate, contributing 
to stable drug entrapment by remaining osmotically inert.37'38,39 This loading 
technology provides substantial stability with negligible drug leakage in circulation, 
while still enabling satisfactory rates of drug release in tissues and malignant 
effusions.40 
Studies in animal tumor models with doxorubicin encapsulated in pegylated 
and other long-circulating liposomes, established the following pharmacologic 
observations41: 
• Increased anti tumor activity of liposomal drug, as compared with optimal doses 
of free drug in various rodent models of syngeneic and human tumors. 
• Increased accumulation of liposomal drug in various transplantable mouse and 
human tumors, compared with free drug. 
• Delayed peak tumor concentration and slow tissue clearance after injection of 
liposomal drug. 
The most valuable pharmacokinetic advantage of the Stealth liposomal delivery 
system is the enhancement of tumor exposure to doxorubicin, as a result of the 
accumulation of Stealth liposomes in tumors, as demonstrated in animal models 
and in some forms of human cancer. When the tissue uptake of PLD was examined 
in a couple of syngeneic mouse tumor models, it was found that the tumor 
drug uptake correlated linearly with dose, while the liver drug uptake showed a 
Liposomal Drug Carriers in Cancer Therapy 449 
saturation profile. In the case of free doxorubicin, liver uptake increased linearly 
with dose, while tumor uptake increased marginally with dose. As a result, the 
delta of tumor drug concentration in favor of PLD was substantially greater at high 
doses.15 These results suggest a passive process of liposomal uptake into tumor, 
with nonsaturable kinetics. 
In preclinical therapeutic studies using a variety of rodent tumors and human 
xenografts in immunodeficient nude mice, PLD was more effective than free doxorubicin 
and other (non-pegylated) formulations of liposomal doxorubicin.11 In a 
few instances, the activity of PLD preparations was matched but not surpassed by 
other non-pegylated, long-circulating preparations of liposomal doxorubicin.42 In 
most of these studies, the improved efficacy of PLD was obtained at milligramequivalent 
doses of the MTD of free doxorubicin, indicating that there was a net 
therapeutic gain per milligram drug, independent of toxicity buffering. An elegant 
study addressed this issue directly by examining the activity of escalating doses of 
PLD and doxorubicin against implants of the mouse 3LL tumor (Lewis lung carcinoma), 
and concluded that the activity of l-2mg/kg Doxil was approximately 
equivalent to 9mg/kg doxorubicin, i.e. a 6-fold enhancement in efficacy.43 Similar 
observations were made in the Ml 09 model, pointing to a 4-fold advantage for 
PLD, compared with free doxorubicin, i.e. a dose of 2.5 mg/kg PLD was at least as 
effective as 10 mg/kg free doxorubicin.15 
There is a large body of preclinical data on other liposome formulations of 
anticancer agents moving into clinical development, or already approved for clinical 
use. In many cases, it is likely that the added value of these formulations has 
not been or will not be sufficient to justify further development, despite positive 
preclinical data. 
6. Liposomal Anth racy dines in the Clinic 
The anthracycline antibiotic doxorubicin has a broad spectrum of antineoplastic 
action and a correspondingly widespread degree of clinical use. In addition to its 
role in the treatment of breast cancer, doxorubicin is indicated in the treatment of 
various cancers of the lymphatic and hematopoyetic systems, gastric carcinoma, 
small-cell cancer of the lung, soft tissue and bone sarcomas, as well as cancer of the 
uterus, ovary, bladder and thyroid. Unfortunately, toxicity often limits the therapeutic 
activity of doxorubicin and may preclude adequate dosing. Other common 
complications of conventional anthracycline therapy include alopecia and doselimiting 
myelosuppression. Most importantly, cardiotoxicity limits the cumulative 
dose of conventional anthracycline that can be given safely.44 
Encapsulation of anthracyclines within liposomes significantly alters their 
pharmacokinetic profiles and promotes selectively high drug concentrations in 
tumors.45 In animal studies, these pharmacologic effects resulted in maintained 
450 Gabizon 
or enhanced anthracycline efficacy and safety in a variety of experimental tumor 
types.46 Improved therapeutic index profiles in clinical trials of liposomal anthracycline 
therapy for Kaposi's sarcoma,47 ovarian cancer,48 breast cancer,49 or multiple 
myeloma50 have been reported. Liposomal anthracycline therapy should be preferred 
when conventional anthracycline therapy is likely to be effective, but the 
required course of treatment would lead to unacceptable risk of toxicity. The relative 
lack of cardiotoxicity with liposomal anthracycline therapy is an important 
asset of the liposomal approach.51 
There are 3 commercial formulations of liposomal anthracyclines that have been 
approved for clinical use: Doxil, Myocet, and Daunoxome. Tables 3 and 4 present 
a comparative list of their tolerated doses and pharmacokinetic parameters41,87-89 
respectively. A summary of their main clinical highlights is presented below. 
6.1. Doxil 
As indicated before, Doxil® (known in Europe as Caelyx®) is a doxorubicin formulation 
in which the drug is encapsulated in PEGylated liposomes (Stealth 
Table 3 Comparative single and cumulative tolerated doses of free and liposomal anthracyclines 
based on their acute/subacute toxicity and cardiac toxicity respectively.1 
Doxorubicin2 Doxil Myocet Daunorubicin Daunoxome 
Maximal 60-75 m g / m 2 50-60 m g / m 2 75mg/m2 90mg/m2 100-120 mg/m2 
Single Does 
Maximal ~450mg/m2 Undetermined ~ 785 m g / m 2 900 m g / m 2 Undetermined 
Cumulative (>650mg/m2) 
Dose3 
Maximal 20-25 m g / 12.5 m g / 25 m g / 30 m g / m 2 / 40 m g / 
Dose m2/week m2/week m2/week week m2/week 
Intensity 
Dose Neutropenia Stomatitis, Neutropenia, Neutropenia Neutropenia, 
Limiting Stomatitis Skin toxicity Stomatitis Mucositis 
Toxicities 
1 Other anthracyclines in clinical use: Epirubicin is an epimer of doxorubicin widely used in breast cancer 
with less cardiotoxicity but also less activity on a per mg basis, owing to faster glucuronidation and faster 
clearance. Its therapeutic index advantage over doxorubicin, if any, is marginal. Idarubicin, an analog 
of daunorubicin, is another clinically approved anthracycline but of less common use. Mitoxantrone, an 
antracenedione, is a drug related to anthracyclines with dose-limiting neutropenia and with cardiotoxic 
potential albeit after longer treatment periods than doxorubicin. It is also approved for clinical use but 
its added value is doubtful since it appears to be somewhat less active than doxorubicin in metastatic 
breast cancer. 
doxorubicin cumulative dose may be substantially increased with co-administration of dexrazoxane, 
a cardioprotective agent. 
3Dose associated with 5% risk of cardiotoxicity. 
Liposomal Drug Carriers in Cancer Therapy 451 
Table 4 Comparative human pharmacokinetics parameters of free and liposomal 
anthracyclines.1 
Distribution Wi 
(hr) 
Terminal tVi 
(hr) 
Clearance 
(mL/hr) 
Volume of 
Distribution (L) 
Dose 
(mg/m2) 
Reference 
Doxorubicin 
Rapid (min) 
42.9 
46,100 
1447 
60 
Swenson 
eta/.87 
Doxil2 
72.9 
ND4 
49 
4.3 
60 
Gabizon 
rffl/.41 
Myocet 
<1.03 
16.4 
5185 
58.3 
60 
Swenson 
etalF 
Daunorubicin 
Rapid (min) 
20.6 
114,750 
-2000 
75 
Riggs88 
Daunoxome 
5.6 
ND4 
408 
3.2 
100 
Bellott 
rffl/.89 
1. To normalize for body surface, values were corrected for an average body surface area of 1.7 m2. 
2. Median values of 4 studies are shown. 
3. Not reported. Extrapolated approximation is shown. 
4. Mono-exponential elimination of liposomal drug from plasma. Terminal clearance phase of 
released drug not detected. 
liposomes), formulated with a hydrogenated (high phase transition temperature) 
PC and cholesterol. Doxil was granted market clearance in 1995 by the US Food 
and Drug Administration (FDA) for use in the treatment of AIDS-related Kaposis 
Sarcoma (KS), in patients with disease that has progressed on prior to combination 
chemotherapy and who are intolerant to such therapy. In 1996, it was granted 
market clearance by the European Union's commission for Proprietary Medicinal 
Products for the same indication. In 1999, Doxil was granted US market clearance 
for use in the treatment of recurrent carcinoma of the ovary in patients with disease 
that is refractory to paclitaxel-and platinum- based chemotherapy regimens. 
In January 2003, the European Commission of the European Union has granted 
centralized marketing authorization to Doxil, as monotherapy for metastatic breast 
cancer in patients who are at increased cardiac risk. In addition, phase II trials have 
been completed in the US and Europe, investigating the safety and efficacy of Doxil 
in multiple myeloma and in other solid tumors including sarcomas, carcinoma of 
head and neck, hepatocellular carcinoma, prostate cancer and the rest. 
Doxil was already recognized 10 years ago as a liposomal doxorubicin formulation 
with unique pharmacokinetics and a dramatic change in the clinical toxicity 
profile. Clinical pharmacokinetic studies have indicated that Doxil prolongs the 
circulation time of doxorubicin dramatically, in agreement with preclinical studies. 
In 1994, we published the results of a pharmacokinetic study in which 15 patients 
452 Gabizon 
were given sequentially the same dose in drug-equivalents of free doxorubicin 
and Doxil.52 A dramatic reduction in the drug clearance and volume of distribution, 
resulting in a 1000-fold increase in AUC with the liposomal formulation, was 
observed. It was also found that nearly all the drug circulating in plasma is in 
liposome-encapsulated form. Metabolites in plasma were undetectable or at very 
low levels. However, they were readily detected in urine 24 hrs or later after injection, 
indicating that the drug has become bioavailable. The following drug distribution 
picture has emerged from this initial study and from more recent ones41: 
1. Drug circulates in plasma for prolonged periods of time (i.e. half-life in the range 
of ~50-80 hours) in liposome encapsulated form. Despite its prolonged presence 
in blood, the drug is not bioavailable, as long as it remains in the interior of a 
circulating liposome. 
2. Most of the injected drug (>95%) is distributed to tissues in liposomeencapsulated 
form. Once in tissues, drug leakage and liposome breakdown 
with or without liposome internalization by cells gradually provides a pool of 
bioavailable drug. Metabolites are formed. 
3. Rate of metabolite production is slower than the rate of renal clearance of metabolites. 
As a result, metabolites do not accumulate in plasma but can be detected 
in urine. 
4. A small fraction of injected drug (<5%) leaks from circulating liposomes and is 
handled as a free drug with fast plasma clearance and rapid metabolism. This 
drug fraction is the source of small amounts of metabolites that can sometimes 
be detected in plasma. 
In 1995, a phase I study of Doxil in patients with solid tumors53 provided clear 
evidence of a major change in the toxicity profile, with muco-cutaneous toxicities 
as the major dose-limiting toxicities. In contrast, myelosuppression, and alopecia 
were minor and cardiotoxicity was conspicuously absent. The maximal tolerated 
dose was established as 60mg/m2, with mucositis being the dose-limiting toxicity. 
It was also found that the optimal dosing interval for retreatment was 4 weeks 
rather than the standard 3-week schedule of doxorubicin. The dose-schedule limiting 
toxicity was a form of skin toxicity known as hand-foot syndrome, also referred 
to as palmar-plantar erythrodysesthesia (PPE), which appears to be related to the 
long half-life of Doxil. Thus, it became well-established that the Doxil liposome formulation 
imparts a significant pharmacokinetic-pharmacodynamic change to the 
drug doxorubicin, unprecedented in magnitude for any intravenous drug delivery 
system. Later on, data gathered from phase II and III studies in metastatic breast 
cancer and recurrent ovarian cancer48,49'54 brought down the recommended dose 
of Doxil to 40-50 mg/m2 once in 4 weeks, i.e. 10-12.5 mg/m2 per week. This dose 
Liposomal Drug Carriers in Cancer Therapy 453 
reduction was needed mainly to prevent skin toxicity resulting from successive 
courses of therapy at a dose intensity of 15 mg/m2/week. 
Kaposi's sarcoma (KS) is a multifocal tumor affecting the skin and sometimes 
the mucosas well known for its extremely high microvascular permeability. Profuse 
extravasation of colloidal gold-labeled Stealth liposomes in a transgenic mouse 
model of KS has been shown.55 In AIDS patients, KS is frequent and has an aggressive 
course. Therefore, this condition was chosen for the initial clinical testing of 
Doxil in Phase II—III studies. Indeed, Doxil as a single agent therapy demonstrated 
a significantly greater efficacy and better safety than standard chemotherapy (i.e. 
combinations of bleomycin and vincristine, with or without doxorubicin) and was 
also effective as second line chemotherapy in pretreated patients. As a result of the 
extreme sensitivity of KS to chemotherapy, a low and relatively subtoxic dose of 
20mg/m2 every 3 weeks is sufficient for effective treatment.47 
Further to the successful application in KS treatment, there are three important 
benchmarks in the clinical research development of Doxil in solid tumors: 
1. Cardiac function in patients receiving Doxil. Evidence of a major risk reduction 
of cardiotoxicity, compared with free doxorubicin historical data. A retrospective 
analysis of patients treated with large cumulative doses of Doxil did not reveal 
any significant cardiac toxicity, despite the fact that some of these patients were 
treated with 3 times as much as the maximal cumulative dose acceptable for free 
doxorubicin.56 Two additional reports focusing on the cardiac biopsies of Doxiltreated 
patients at high cumulative doses confirmed the cardiac safety of Doxil.57,58 
2. Phase III study in recurrent ovarian cancer. Significant increase in median survival 
with improved safety profile in the Doxil patient group versus the topotecan-treated 
patient group, where topotecan was the former standard therapy in this condition.59 
The use of Doxil was particularly beneficial in "platinum-sensitive" patients (i.e. 
patients in whom tumor recurrence occurred more than 6 months after the discontinuation 
of platinum-based front-line therapy). In this subgroup, the median 
survival of Doxil-treated patients was 107.9 weeks, compared with 70.1 weeks for 
topotecan-treated patients, a difference of ~9 months, equivalent to a 54% increase 
in survival. As a result, Doxil has become the standard therapy for recurrent ovarian 
cancer. 
3. Phase III study in metastatic breast cancer. Equivalent anti tumor activity and 
reduced cardiotoxicity. In this study, it was found that treatment with Doxil 
50mg/m2 every 4 weeks (dose intensity = 12.5 mg/m2/week) had equivalent 
efficacy to free doxorubicin of 60 mg/m2 every 3 weeks (dose intensity 
20 mg/m2/week), despite the lower dose intensity of the former.60 In addition, cardiotoxicity, 
as well as alopecia, were dramatically reduced in Doxil-treated patients. 
Myelosuppression and nausea were also milder in Doxil-treated patients. However, 
454 Gabizon 
as seen in phase I—II studies, skin toxicity was prominent with Doxil and almost 
absent with doxorubicin treatment. 
6.2. Myocet 
Myocet™ (liposome encapsulated doxorubicin citrate complex) is a non-pegylated 
formulation of liposomal doxorubicin that has been approved for the treatment of 
metastatic breast cancer in Europe, but not in the United States. Myocet lipid composition 
consists of a fluid phase (low phase transition temperature) PC, as well as 
cholesterol with doxorubicin entrapped in the water phase. A unique feature of this 
formulation is that the drug is loaded into liposomes using a 3-vial kit, just prior 
to administration in the hospital pharmacy. The strategy employed is to decrease 
toxicity in a way that a net gain of therapeutic index is obtained. In the pivotal 
phase I trial of Myocet,61 the maximum tolerated dose was between 75-90 mg/m2 , 
given on day 1 or split in 3 consecutive daily doses of each 3-week cycle. The major 
dose-limiting toxicity was neutropenia. G-CSF (granulocyte-colony stimulating factor) 
administration may increase the maximum tolerated dose, when higher doses 
of Myocet are desired by reducing the incidence of dose-limiting neutropenia.62 
Phase III studies in metastatic breast cancer, comparing Myocet to free doxorubicin, 
have shown similar anti tumor activity with significantly lesser cardiotoxicity for 
Myocet.63,64 However, a study of high dose Myocet with G-CSF in patients with 
advanced breast cancer resulted in a disturbingly high incidence of cardiotoxicity 
(38%),65 after a median cumulative dose of 405 mg/m2 (range of 135-1065 mg/m2), 
suggesting that the Myocet margin of cardiotoxicity gain over doxorubicin is 
limited. 
As described above, the maximum tolerated doses of Doxil and Myocet differ as 
a result of their different formulations. Consequently, a more intense dose schedule 
is recommended for Myocet (75 mg/m2 every 3 weeks) than for Doxil (50 mg/m2 
every 4 weeks). The pharmacokinetics of Myocet points to a small change in clearance 
and volume of distribution, when compared with free doxorubicin,66 suggesting 
that, in contrast to Doxil, Myocet liposomes are cleared rapidly from circulation 
and their drug content also leaks substantially. 
6.3. Daunoxome 
Daunoxome® (daunorubicin citrate liposome injection), is approved for use in 
patients with advanced HIV-associated Kaposi's sarcoma at a recommended dose of 
40 mg/m2 every 2 weeks.67 Daunoxome lipid composition consists of a solid phase 
(high phase transition temperature) PC, as well as cholesterol with daunorubicin 
encapsulated in the water phase. The maximum tolerated dose of Daunoxome was 
Liposomal Drug Carriers in Cancer Therapy 455 
evaluated in a phase I trial of 32 patients with solid tumors, using Daunoxome doses 
that were escalated in steps from 10 to 120mg/m2.68 Dose-limiting neutropenia 
occurred in all patients who received 120 mg/m2 . The maximum tolerated dose of 
Daunoxome was established as 100-120 mg/m2 in patients with solid tumors every 
three weeks. A more recent study with Daunoxome in breast cancer patients confirmed 
this dose level as MTD.69 Two studies in patients with acute leukemia administered 
high dose Daunoxome in 3-day courses, up to a total dose of 450 mg/m2.70'71 
Mucositis was the dose-limiting toxicity in these high dose Daunoxome studies. In 
one of the studies, 2 of 28 patients experienced fatal cardiotoxicity.71 In a phase I 
study of 48 children with relapsed or resistant solid tumors,72 the trial was prematuredly 
discontinued due to evidence of cumulative cardiotoxicity, including two 
deaths after 4 courses of Daunoxome treatment. In summary, neutropenia typically 
limits the maximum dose of Daunoxome in solid tumor patients, and mucositis 
limits the maximum dose that can be given in a single cycle to leukemia patients. 
Cardiotoxicity was reported in some patients, and appears especially problematic 
in children. 
The pharmacokinetics of Daunoxome points to a major retardation of clearance 
and a small volume of distribution, when compared with free daunorubicin.67 The 
distribution half-life of Daunoxome (7-8 h) is nearly 10 times shorter than that of 
Doxil, suggesting that the former is cleared faster by the RES, in agreement with the 
differences observed between pegylated and nonpegylated liposomes in preclinical 
studies. However, in contrast to Myocet, the Daunoxome pharmacokinetics data do 
not point at any significant drug leakage in circulation. 
7. Clinical Development of Other Liposome-entrapped 
Cytotoxic Agents 
Dose-finding safety studies have been performed with several other liposomal 
anthracyclines, including a cardiolipin-based liposome formulation of 
doxorubicin,73 doxorubicin entrapped in negatively charged phosphatidylglycerol 
liposomes,74 liposomal annamycin,75 a pegylated form of liposomal doxorubicin 
manufactured only in Taiwan (PEG-distearoylPC-cholesterol),76 and an 
immunoliposome-encapsulated form of doxorubicin (MCC-465), targeted with an 
antibody reacting with human gastric cancer.77 Some of these formulations have 
not progressed any further to phase II—III studies. For those formulations still under 
investigation, further clinical trials are needed to establish their efficacy. 
The most advanced compounds currently being developed in liposomes 
belong to two families of cytotoxic drugs: Camptothecin analogs with topoisomerase 
I inhibitory properties,78,79 and Vinca alkaloids.80 Both are cell cycle phasespecific 
cytotoxic drugs. Their anti-tumor activity tends to increase with liposome 
456 Gabizon 
encapsulation, as we extend time of exposure by exploiting the liposomal slow 
release features. An additional advantage for camptothecin analogs is the fact that 
their activity is better maintained in the lactone configuration which is stable in an 
acid environment, as is the case of liposomes with remote drug-loading techniques 
based on proton gradients. At this time, none of these compounds have yet been 
approved by a regulatory agency after going through phase III studies. 
It should be noted that the development of liposomal formulations of cytotoxic 
agents has often failed for different reasons. The formulation of a liposomal anticancer 
agent is a complex process with at least three distinct variables that may 
affect the outcome and the risk of failure: the choice of the liposome carrier, the 
choice of the drug, and the method of drug encapsulation. One example is the formulation 
of cisplatin in pegylated liposomes known as SPI-77. These liposomes 
have long-circulating characteristics, retain the drug in plasma exceedingly well, 
reduce drug toxicity, and produce a high and long-lasting accumulation of drug 
in implanted tumors.81 However, the anti-tumor activity is reduced by comparison 
to free drug in preclinical studies.81 In clinical studies, SPI-77 was inactive and 
showed no dose-limiting toxicities, even at doses greater than 2-fold the MTD of 
free cisplatin.82'83 It appeared that cisplatin release from these liposomes was minimal 
in in vitro81 as well as in vivo systems, as indicated by the low occurrence of 
DNA adducts resulting in a major reduction of bioavailability.84 As a result, the 
development of SPI-77 was discontinued at the Phase I clinical level. 
8. The Future of Liposomal Nanocarriers 
Although progress in the understanding of cancer biology at the molecular level 
will undoubtedly lead to more new drugs with exquisite selectivity, some of these 
agents may need an efficient system for in vivo delivery to yield optimal results. 
In addition, the use of broad-spectrum cytotoxic drugs will remain as a major tool 
in cancer therapy for decades to come, and, for cytotoxic drugs, various delivery 
systems may have a beneficial effect on the therapeutic index as already shown. 
Liposomes remain one of the most attractive platforms for systemic drug delivery, 
and an increased sophistication of these systems would be expected. The most 
immediate improvement is the coupling of a ligand to the surface of the liposome 
that will target the vesicles to a specific cell-surface receptor, followed in most cases 
by internalization and intracellular delivery of the liposome drug cargo. Examples 
in this direction are the targeting of Doxil to Her 2 expressing or folate-receptor, 
expressing cancer cells using a specific anti-Her 2 single chain Fv respectively or a 
folate conjugate anchored to the liposome surface.30,85 Another example is the tumor 
vascular targeting by endothelium-specific peptides associated to liposomes.86 A 
Liposomal Drug Carriers in Cancer Therapy 457 
major advantage of targeted liposomal nanocarriers over ligand-drug bioconjugates 
is the delivery-amplifying effect of the former, which may be able to provide to the 
target cell, a ratio of 1000 drug molecules per single ligand-receptor interaction. 
Other future avenues that can be exploited using the liposome platform and 
have a definite preclinical basis include: 
1. Association with a reporter, i.e. an imaging or tracing element, that will provide 
the possibility of tracking the fate of the liposome in vivo or even the occurrence 
of a pharmacological effect such as apoptosis. 
2. Co-delivery of synergistic agents. The liposome platform offers the possibility of 
co-delivery in space and time of two drugs with different pharmacokinetics and 
biodistribution patterns, thus enabling the optimal exploitation of synergistic 
properties. 
3. Association with a bio-responsive element, i.e. a temperature-sensitive or pHsensitive 
component that could destabilize the liposome and drive a burst release 
of drug. 
The fact that liposome technology has matured into an acceptable pharmaceutical 
technology, and the promising contributions of liposomes to sophisticated drug 
delivery methods, augur that liposome carriers are to remain in cancer therapy for 
the foreseeable future. 
References 
1. Cockshott ID (2000) Clinical pharmacokinetics of goserelin. Clin Pharmacokinet 39: 
27^:8. 
2. Ibrahim NK, Desai N, Legha S, et al. (2002) Phase I and pharmacokinetic study of ABI- 
007, a Cremophor-free, protein-stabilized, nanoparticle formulation of paclitaxel. Clin 
Cancer Res 8:1038-1044. 
3. Duncan R (2003) The dawning era of polymer therapeutics. Nat Rev Drug Discov 1: 
347-360. 
4. Moghimi SM, Hunter AC and Murray JC (2005) Nanomedicine: Current status and 
future prospects. FASEB } 19:311-330. 
5. Hanahan D and Weinberg RA (2000) The hallmarks of cancer. Cell 100:57-70. 
6. Van Den Bossche B and Van de Wiele C (2004) Receptor imaging in oncology by means 
of nuclear medicine: Current status. / Clin Oncol 22:3593-3607. 
7. Kastan MB (1997) Molecular biology of cancer: The cell cycle, in Cancer: Principles 
& Practice of Oncology, 5th ed., DeVita Jr. VT, Hellman S and Rosenberg SA (eds.), 
Lippincott-Raven: Philadelphia, pp. 121-134. 
8. Gabizon A (2001) Pegylated liposomal doxorubicin: Metamorphosis of an old drug into 
a new form of chemotherapy. Cancer Invest 19:424^436. 
458 Gabizon 
9. Jain RK (2001) Delivery of molecular medicine to solid tumors: Lessons from in vivo 
imaging of gene expression and function. / Control Rel 74:7-25. 
10. Papahadjopoulos D, Allen T, Gabizon A, et al. (1991) Sterically stabilized liposomes: 
Improvements in pharmacokinetics and anti tumor therapeutic efficacy. Proc Natl Acad 
Sci USA 88:11460-11464. 
11. Gabizon A and Martin F (1997) Polyethylene-glycol-coated (Pegylated) Liposomal 
Doxorubicin- Rationale for use in solid tumors. Drugs 54(Suppl. 4):15-21. 
12. Swabb EA, Wei J and Gullino PM (1974) Diffusion and convection in normal and neoplastic 
tissues. Cancer Res 34:2814-2822. 
13. Harrington KJ, Rowlinson-Busza G, Syrigos KN, et al. (2000) Influence of tumor size 
on uptake of (lll)ln-DTPA-labelled pegylated liposomes in a human tumor xenograft 
model. Br f Cancer 83:684-688. 
14. Safra T, Groshen S, Jeffers S, et al. (2001) Treatment of patients with ovarian carcinoma 
with pegylated liposomal doxorubicin: Analysis of toxicities and predictors of outcome. 
Cancer 91:90-100. 
15. Gabizon A, Tzemach D, Mak L, et al. (2002) Dose dependency of pharmacokinetics 
and therapeutic efficacy of pegylated liposomal doxorubicin (Doxil) in murine models. 
/ Drug Targ 10:539-548. 
16. Gregoriadis G and Ryman BE (1971) Liposomes as carriers of enzymes or drugs: A new 
approach to the treatment of storage diseases. Biochem } 124:58P 
17. Allen TM and Chonn A (1987) Large unilamellar liposomes with low uptake into the 
reticuloendothelial system. FEBS Lett 223:42-46. 
18. Gabizon A and Papahadjopoulos D (1988) Liposome formulations with prolonged circulation 
time in blood and enhanced uptake by tumors. Proc Natl Acad Sci USA 85: 
6949-6953. 
19. Gabizon A, Shiota R and Papahadjopoulos D (1989) Pharmacokinetics and tissue distribution 
of doxorubicin encapsulated in stable liposomes with long circulation times. 
J Natl Cancer Inst 81:1484-1488. 
20. Woodle MC and Lasic DD (1992) Sterically stabilized liposomes. Biochim Biophys Acta 
1113:171-199. 
21. Klibanov AL, Maruyama K, Torchilin VP, et al. (1990) Amphipathic polyethyleneglycols 
effectively prolong the circulation time of liposomes. FEBS Lett 268:235-237. 
22. Allen TM, Hansen C, Martin F, et al. (1991) Liposomes containing synthetic lipid 
derivatives of poly (ethylene glycol) show prolonged circulation half-lives in vivo. Biochim 
Biophys Acta 1066:29-36. 
23. Lasic DD and Martin FJ (eds.) Stealth Liposomes. CRC Press: Boca Raton, 1995. 
24. Huang SK, Lee K-D, Hong K, et al. (1992) Microscopic localization of sterically stabilized 
liposomes in colon carcinoma-bearing mice. Cancer Res 52:5135-5143. 
25. Yuan F, Leunig M, Huang SK, et al. (1994) Microvascular permeability and interstitial 
penetration of sterically stabilized (stealth) liposomes in a human tumor xenograft. 
Cancer Res 54:3352-3356. 
26. Gabizon AA (1992) Selective tumor localization and improved therapeutic index of 
anthracyclines encapsulated in long-circulating liposomes. Cancer Res 52:891-896. 
Liposomal Drug Carriers in Cancer Therapy 459 
27. Bally MB, Masin D, Nayar R, et al. (1994) Transfer of liposomal drug carriers from the 
blood to the peritoneal cavity of normal and ascitic tumor-bearing mice. Cancer Chemother 
Pharmacol 34:137-146. 
28. Maeda H (2001) Enhanced permeability and retention (EPR) effect in tumor vasculature: 
The key role of tumor-selective macromolecular drug targeting. Adv Enzyme Regul 
41:189-107. 
29. Goren D, Horowitz AT, Zalipsky S, et al. (1996) Targeting of stealth liposomes to erbB2 
(Her/2) receptor: In vitro and in vivo studies. Br } Cancer 74:1749-1756. 
30. Gabizon A, Shmeeda H, Horowitz AT, et al. (2004) Tumor cell targeting of liposomeentrapped 
drugs with phospholipid-anchored folic acid-PEG conjugates. Adv Drug Del 
Rev 56:1177-1192. 
31. Stohrer M, Boucher Y, Stangassinger M, et al. (2000) Oncotic pressure in solid tumors is 
elevated. Cancer Res 60:4251-4255. 
32. Harrington KJ, Mohammadtaghi S, Uster PS, et al. (2001) Effective targeting of solid 
tumors in patients with locally advanced cancers by radiolabeled Pegylated liposomes. 
Clin Cancer Res 7:243-254. 
33. Symon Z, Peyser A, Tzemach D, et al. (1999) Selective delivery of doxorubicin to patients 
with breast carcinoma metastases by stealth liposomes. Cancer 86:72-78. 
34. Northfelt DW, Martin FJ, Working P, et al. (1996) Doxorubicin encapsulated in liposomes 
containing surface-bound polyethylene glycol: Pharmacokinetics, tumor localization, 
and safety in patients with AIDS-related Kaposi's sarcoma. / Clin Pharmacol 36:55-63. 
35. Young RC, Ozols RF and Myers CE (1981) The anthracycline antineoplastic drugs. 
N Engl J Med 305:139-153. 
36. Forssen EA and Tokes ZA (1979) In vitro and in vivo studies with adriamycin liposomes. 
Biochem Biophys Res Commun 91:1295-1301. 
37. Lasic DD, Frederik PM, Stuart MCA, et al. (1992) Gelation of liposome interior. A novel 
method for drug encapsulation. FEBS Lett 312:255-258. 
38. Haran G, Cohen LK, Bar Y, et al. (1993) Transmembrane ammonium sulfate gradients in 
liposomes produce efficient and stable entrapment of amphipathic weak bases. Biochim 
Biophys Acta 1151:201-215. 
39. Lasic DD, Ceh B, Stuart MC, et al. (1995) Transmembrane gradient driven phase transitions 
within vesicles: Lessons for drug delivery. Biochim Biophys Acta 1239:145-156. 
40. Gabizon AA (1995) Liposome circulation time and tumor targeting: Implications for 
cancer chemotherapy. Adv Drug Del Rev 16:285-294 
41. Gabizon A, Shmeeda H and Barenholz Y (2003) Pharmacokinetics of pegylated liposomal 
doxorubicin: Review of animal and human studies. Clin Pharmacokinetics 42:419^36. 
42. Gabizon A, Chemla M, Tzemach D, et al. (1996) Liposome longevity and stability in circulation: 
Effects on the in vivo delivery to tumors and therapeutic efficacy of encapsulated 
anthracyclines. / Drug Targ 3:391-398. 
43. Colbern GT, Hiller AJ, Musterer RS, et al. (1999) Significant increase in antitumor potency 
of doxorubicin HCI by its encapsulation in pegylated liposomes. / Lipos Res 9:523-538. 
44. Hortobagyi GN (1997) Anthracyclines in the treatment of cancer: An overview. Drugs 
54(suppl 4): 1-7. 
460 Cabizon 
45. Allen TM and Martin F (2004) Advantages of liposomal delivery systems for anthracyclines. 
Semin Oncol 31(suppl 13):5-15 
46. Vail DM, Amantea MA, Colbern GT, et al. (2004) Pegylated liposomal anthracycline 
formulations: Proof of principle using preclinical animal models and pharmacokinetics. 
Semin Oncol 31(suppl 13):16-35. 
47. Krown SE, Northfelt DW, Osoba D, et al. (2004) Use of liposomal anthracyclines in 
Kaposi's sarcoma. Semin Oncol 31(suppl 13):36-52. 
48. Markman M, Gordon AN, McGuire WP, et al. (2004) Liposomal anthracycline treatment 
for ovarian cancer. Semin Oncol 31(suppl 13):91-105. 
49. Robert NJ, Vogel CL, Henderson IC, et al. (2004) The role of the liposomal anthracyclines 
and other systemic therapies in the management of advanced breast cancer. Semin Oncol 
31(suppl 13):106-146. 
50. Hussein MA and Anderson KC (2004) Role of liposomal anthracyclines in the treatment 
of multiple myeloma. Semin Oncol 31(suppl 13):147-160. 
51. Ewer MS, Martin FT, Henderson IC, et al. (2004) Cardiac safety of liposomal anthracyclines. 
Semin Oncol 31(suppl 13):161—181. 
52. Gabizon A, Catane R, Uziely B, et al. (1994) Prolonged circulation time and enhanced 
accumulation in malignant exudates of doxorubicin encapsulated in poly ethylene-glycol 
coated liposomes. Cancer Res 54:987-992. 
53. Uziely B, Jeffers S, Isacson R, et al. (1995) Liposomal doxorubicin: Antitumor activity and 
unique toxicities during two complementary phase I studies. / Clin Oncol 13:1777-1785. 
54. Alberts DS, Muggia FM, Carmichael J, et al. (2004) Efficacy and safety of liposomal 
anthracyclines in phase I/II clinical trials. Semin Oncol 31(Suppl 13):53-90. 
55. Huang SK, Martin FT, Jay G, et al. (1993) Extravasation and transcytosis of liposomes in 
Kaposi's sarcoma-like dermal lesions of transgenic mice bearing the HIV tat gene. Am } 
Pathol 143:10-14. 
56. Safra T, Muggia M, Jeffers S, et al. (2000) Pegylated Liposomal Doxorubicin: Reduced 
clinical cardiotoxicity in patients reaching or exceeding 500 mg/m2 cumulative doses of 
Doxil. Ann Oncol 11:1029-1033. 
57. Berry G, Billingham M, Alderman E, et al. (1998) The use of cardiac biopsy to demonstrate 
reduced cardiotoxicity in AIDS Kaposi's sarcoma patients treated with pegylated 
liposomal doxorubicin. Ann Oncol 9:711-716. 
58. Gabizon AA, Lyass O, Berry GJ, et al. (2004) Cardiac safety of pegylated liposomal doxorubicin 
(doxil®/caelyx®) demonstrated by endomyocardial biopsy in patients with 
advanced malignancies. Cancer Invest 22:663-669. 
59. Gordon AN, Tonda M, Sun S, et al. (2004) Long-term survival advantage for women 
treated with pegylated liposomal doxorubicin compared with topotecan in a phase 3 
randomized study of recurrent and refractory epithelial ovarian cancer. Gynecologic Oncol 
95:1-8. 
60. O'Brien ME, Wigler N, Inbar M, et al. (2004) Reduced cardiotoxicity and comparable efficacy 
in a phase III trial of pegylated liposomal doxorubicin HC1 (Caelyx/Doxil) versus 
conventional doxorubicin for first-line treatment of metastatic breast cancer. Ann Oncol 
15:440^149. 
Liposomal Drug Carriers in Cancer Therapy 461 
61. Cowens JW, Creaven PJ, Greco WR, et al. (1993) Initial clinical (phase I) trial of TLC D-99 
(doxorubicin encapsulated in liposomes). Cancer Res 53:2796-2802. 
62. Shapiro CL, Ervin T, Welles L, et al. (1999) Phase II trial of high-dose liposomeencapsulated 
doxorubicin with granulocyte colony-stimulating factor in metastatic 
breast cancer. TLC D-99 Study Group. / Clin Oncol 17:1435-1441. 
63. Harris L, Batist G, Belt R, et al. (2002) Liposome-encapsulated doxorubicin compared 
with conventional doxorubicin in a randomized multicenter trial as first-line therapy of 
metastatic breast carcinoma. Cancer 94:25-36. 
64. Batist G, Ramakrishnan G, Rao CS, et al. (2001) Reduced cardiotoxicity and preserved 
antitumor efficacy of liposome-encapsulated doxorubicin and cyclophosphamide compared 
with conventional doxorubicin and cyclophosphamide in a randomized, multicenter 
trial of metastatic breast cancer. / Clin Oncol 19:1444-1454. 
65. Casper ES, Schwartz GK, Sugarman A, et al. (1997) Phase I trial of dose-intense liposomeencapsulated 
doxorubicin in patients with advanced sarcoma. / Clin Oncol 15:2111-2117. 
66. Conley BA, Egorin MJ, Whitacre MY, et al. (1993) Phase I and pharmacokinetic trial of 
liposome-encapsulated doxorubicin. Cancer Chemother Pharmacol 33:107-112. 
67. Gill PS, Espina BM, Muggia F, et al. (1995) Phase I/II clinical and pharmacokinetic evaluation 
of liposomal daunorubicin. / Clin Oncol 13:996-1003. 
68. Guaglianone P, Chan K, DelaFlor-Weiss E, et al. (2004) Phase I and pharmacologic study 
of liposomal daunorubicin (DaunoXome). Invest New Drugs 12:103-110. 
69. O'Byrne KJ, Thomas AL, Sharma RA, et al. (2002) A phase I dose-escalating study of 
DaunoXome, liposomal daunorubicin, in metastatic breast cancer. Br } Cancer 87:15-20. 
70. Cortes J, O'Brien S, Estey E, et al. (1999) Phase I study of liposomal daunorubicin in 
patients with acute leukemia. Invest New Drugs 17:81-87. 
71. Fassas A, Buffels R, Anagnostopoulos A, et al. (2002) Safety and early efficacy assessment 
of liposomal daunorubicin (DaunoXome) in adults with refractory or relapsed acute 
myeloblasts leukaemia: A phase I—II study. Br } Haematol 116:308-315. 
72. Lowis S, Lewis I, Elsworth A, et al. (2002) Cardiac toxicity may limit the usefulness of liposomal 
daunorubicin (DaunoXome): Results of a phase I study in children with relapsed 
or resistant tumours — a UKCCSG/SFOP study. Proc Am Soc Clin Oncol 21:Abstract 435. 
73. Rahman A, Treat J, Roh JK, et al. (1990) A phase I clinical trial and pharmacokinetic 
evaluation of liposome-encapsulated doxorubicin. / Clin Oncol 8:1093-1100. 
74. Gabizon A, Peretz T, Sulkes A, et al. (1989) Systemic administration of doxorubicincontaining 
liposomes in cancer patients: A phase I study. Eur } Cancer Clin Oncol 25: 
1795-1803. 
75. Booser DJ, Perez-Soler R, Cossum P, et al. (2000) Phase I study of liposomal annamycin. 
Cancer Chemother Pharmacol 46:427-432 
76. Hong RL and Tseng YL (2001) Phase I and pharmacokinetic study of a stable, 
polyethylene-glycolated liposomal doxorubicin in patients with solid tumors: The relation 
between pharmacokinetic property and toxicity. Cancer 91:1826-1833. 
77. Matsumura Y, Gotoh M, Muro K, et al. (2004) Phase I and pharmacokinetic study of 
MCC-465, a doxorubicin (DXR) encapsulated in PEG immunoliposome, in patients with 
metastatic stomach cancer. Ann Oncol 15:517-525. 
462 Gabizon 
78. Kehrer DF, Bos AM, Verweij J, et al. (2002) Phase I and pharmacologic study of liposomal 
lurtotecan, NX 211: Urinary excretion predicts hematologic toxicity. / Clin Oncol 20: 
1222-1231. 
79. Colbern GT, Dykes DJ, Engbers C, et al. (1998) Encapsulation of the topoisomerase I 
inhibitor GL147211C in pegylated (STEALTH) liposomes: Pharmacokinetics and antitumor 
activity in HT29 colon tumor xenografts. Clin Cancer Res 4:3077-3082. 
80. Gelmon KA,Tolcher A, Diab AR, et al. (1999) Phase I study of liposomal vincristine. 
/ Clin Oncol 17:697-705. 
81. Bandak S, Goren D, Horowitz AT, et al. (1999) Pharmacological studies of cisplatin encapsulated 
in long-circulating liposomes in mouse tumor models. Anti-Cancer Drugs 10: 
911-920. 
82. Veal GJ, Griffin MJ, Price E, et al. (2001) A phase I study in pediatric patients, to evaluate 
the safety and pharmacokinetics of SPI-77, a liposome encapsulated formulation of 
cisplatin. Br] Cancer 84:1029-1035. 
83. Harrington KJ, Lewanski CR, Northcote AD, et al. (2001) Phase I—II study of pegylated 
liposomal cisplatin (SPI-077) in patient with inoperable head and neck cancer. Ann Oncol 
12:493^96. 
84. Meerum Terwogt JM, Groenewegen G, Pluim D,etal. (2002) Phase I and pharmacokinetic 
study of SPI-77, a liposomal encapsulated dosage form of cisplatin. Cancer Chemother 
Pharmacol 49:201-210. 
85. Park JW, Hong K, Kirpotin DB, et al. (2002) Anti-HER2 immunoliposomes: Enhanced 
efficacy attributable to targeted delivery. Clin Cancer Res 8:1172-1181. 
86. Pastorino F, Brignole C, Marimpietri D, et al. (2003) Vascular damage and anti-angiogenic 
effects of tumor vessel-targeted liposomal chemotherapy. Cancer Res 63:7400-7409. 
87. Swenson CE, Bolcsak LE, Batist G, et al. (2003) Pharmacokinetics of doxorubicin administered 
i.v. as Myocet (TLC D-99; liposome-encapsulated doxorubicin citrate) compared 
with conventional doxorubicin when given in combination with cyclophosphamide in 
patients with metastatic breast cancer. Anti-Cancer Drugs 14:239-246. 
88. Riggs CE (1984) Clinical pharmacology of daunorubicin in patients with acute leukemia. 
Semin Oncol 11(4 Suppl 3):2-ll. 
89. Bellott R, Auvrignon A, Leblanc T, et al. (2001) Pharmacokinetics of liposomal 
daunorubicin (DaunoXome) during a phase I—II study in children with relapsed acute 
lymphoblastic leukemia. Cancer Chemother Pharmacol 47:15-21. 
21 
Nanoparticulate Drug Delivery to the 
Reticuloendothelial System and to 
Associated Disorders 
Mukul Kumar Basu and Sanchaita Lala 
1. Introduction 
Physicochemical methods for site-specific delivery of drugs in the form of polymeric 
and colloidal particulate carriers e.g. liposomes, micelles and nanoparticles, 
have been of great interest to researchers recently. Nanoparticles, particularly polymeric 
nanoparticles, have been investigated since the late 1970s as an alternative 
to liposomes which suffered from inherent problems such as low encapsulation, 
rapid leakage, poor stability and production difficulties. Nanoparticles are particles 
ranging from 10 nm to 1000 nm in diameter and are the collective names for 
nanospheres and nanocapsules. Nanospheres are solid particles which can be used 
as drug carriers, where the active principles can be dispersed in the polymeric 
matrix or adsorbed on their surfaces or both. Nanocapsules have a polymeric shell 
with an inner liquid core and the active substances can be incorporated within 
the core or loaded on the surface by physical adsorption or by chemical bonding. 
Nanoparticles can be administered by a variety of routes, depending upon the 
desired therapeutic outcome and can even be used for vaccine administration and 
diagnostic imaging. 
Drug release from nanoparticles is a very important factor for developing successful 
formulations. For achieving this, various targeted and controlled release 
463 
464 Basu & Lala 
drug delivery systems have been developed. Controlled release is one of the basic 
modes of drug delivery with the objective of releasing a drug into a patient's body 
at a predetermined rate, or at specific times, or with specific release profiles. Release 
profiles of the drugs from nanoparticles depend on the nature of the delivery systems. 
Such systems often use synthetic polymers as the carriers for the drug. 
These particulate carrier systems can release the drug by (a) polymer degradation 
or chemical cleavage of the drug from the polymer, (b) swelling of the polymers 
and releasing the drug entrapped within them, (c) osmotic pressure effects creating 
pores and releasing the drugs, which can get dispersed within the polymeric 
network of the nanoparticles, and (d) simple diffusion methods. They represent 
an interesting carrier system for the specific enrichment in macrophage containing 
organs like the liver and spleen. Thus, injectable nanoparticle carriers have important 
potential application in site-specific drug delivery. 
2. Reticuloendothelial System and Associated Disorders 
The reticuloendothelial system (RES) represents a group of cells having the ability 
to take up and sequester inert particles and vital dyes. This includes macrophages 
and macrophage precursors, specialized endothelial cells lining the sinusoids of 
the liver, spleen and bone marrow, and the reticular cells of the lymphatic tissue 
(macrophages) and of the bone marrow (fibroblasts). Thus, the reticuloendothelial 
system or the mononuclear phagocyte system encompasses a range of cells capable 
of phagocytosis i.e. macrophages and monocytes. They are either freely circulating 
within the blood or fixed to various connective tissues. Examples of the sites of fixed 
cells include pulmonary alveoli, liver sinusoids, skin, spleen and joints. The RES 
primarily functions to (a) remove the senescent cells from circulation and (b) provide 
phagocytic cells for both inflammatory and immune responses. 
There are several RES associated disorders involving both macrophages and 
monocytes. These diseases could be of two types — infectious and non-infectious. 
Among the infectious diseases, pulmonary tuberculosis, typhoid fever, leishmaniasis, 
trypanosomiasis and acquired immune deficiency syndrome (AIDS) are worth 
mentioning. Among the non-infectious type, ulcerative colitis, collagen disease, 
Hodgkin's disease and Gaucher's disease are equally common. 
3. Uptake of Nanoparticles by the 
Reticuloendothelial System 
3.1. Sites of uptake 
It has been observed that nanoparticles, like other colloidal drug delivery systems 
e.g. liposomes, niosomes, microparticles etc., on intravenous injections, are rapidly 
Nanoparticulate Delivery to Reticuloendothelial System 465 
sequestered and retained by the organs comprising of the reticuloendothelial systems 
(RES), mainly the liver, spleen and the bone marrow. Thus, targeting of the 
nanoparticles to the RES is much simpler than to any other organ. In the liver, the 
particles are mainly retained by the scavenging periportal and midzonal Kupffer 
cells, while the hepatocytes and liver endothelial cells may play a secondary role 
under special pathophysiological conditions or for special physico-chemical characteristics 
of particles. In the spleen, the marginal zone and red pulp macrophages 
are the major scavengers, while peritoneal macrophages and dendritic cells have 
a minor contribution. In the case of bone marrow, the sequestering mechanism is 
rather complex and appears to be species specific. Briefly, it encompasses two routes, 
a transcellular route, mediated through the diaphragmed fenestrate of the endothelial 
walls, and an intercellular route, involving the formation of bristle-coated pits 
containing matter on the internal surface of the endothelium. The particles that 
reach the bone marrow parenchyma are engulfed by stromal or hematopoietic 
macrophages. In guineapig, the venous endothelium has similar phagocytic properties 
to the sinusoidal endothelium, while in some other species such as rabbit 
and marmoset, certain macrophage subpopulations named perisinal phagocytes, 
which penetrate the endothelium sending cytoplasmic processes into the lumen, 
are also involved.12 
3.2. Mechanism of uptake 
Two major mechanisms are involved in the sequestering of nanoparticles and 
other colloidal particles by macrophages. These are (a) opsonophagocytosis and 
(b) opsonic-independent recognition. 
(a) Opsonophagocytosis 
Opsonophagocytosis and opsonorecognition of particles by macrophage receptors 
are greatly enhanced when the particle surface is coated with certain protein 
ligands. Such ligands are known as "opsonins" and the process of their adsorption 
on the particle surface is termed "opsonization". Opsonins act as a ligand 
on the particle surface and facilitate their recognition and initial attachment by 
the phagocyte receptors. Taking liposomes as a model system, the major classes 
of opsonins have been found to include various subclasses of immunoglobulins 
(e.g. IgM, IgGl and IgG3 in humans), certain components of the complement system 
(e.g. C3b, iC3b, Clq), fibronectin, lipopolysaccharide binding protein and pentraxins 
(e.g. C-reactive proteins and serum amyloid P-components). More recently, 
it was shown that both thrombospondin and Von Willebrand factor can also act 
as opsonins and trigger phagocytosis of sulfa tide-rich or coated particles. Their 
466 Basu & Lata 
modes of interaction have been reviewed in detail.2 Some opsonic entities such as 
the tetrapeptide tuftsin, unlike other opsonins, bind to the phagocytic cell rather 
than to the particle and enhance the phagocytic ingestion two or three-fold.3 Of 
these, IgG along with complement C3b and albumin are seen to play the most 
important roles. These opsonins, bound to the particle surface, form a bridge 
between the particles and the macrophages, facilitating phagocytosis of the particles. 
The Fc receptor plays an important role in the clearance of IgG-opsonized 
particles. 
The complement system provides the first line of defense against foreign 
microbes and particles, ensuring their cytolytic and/or phagocytic clearance. As 
in the case of liposomes, activation of the complement pathway may occur if the 
opsonic components C3b and iC3b are deposited on nanoparticles. This may be 
antibody-mediated, or nanoparticles may also activate complement through nonantibody 
mediated mechanism via the classical and alterative pathways. However, 
it is known that non-covalent linkage of C3b and iC3b to a particle surface cannot 
promote phagocytosis by macrophages. Rather, the covalent bond between 
the reactive glutamyl residue of the C3 thiol ester and the constituents of the 
particles surface is the central mechanism of opsonization and the mediator of 
phagocytic recognition.4 In addition, the covalently attached C3b or iC3b must be 
accessible to their corresponding receptors on phagocytic cell surface. Moreover, 
the complement receptors must be in a functional state. Kupffer cells in humans 
and rats have been found to contain receptors for C3b (CR1) and iC3b (CR3 and 
CR4), and may play an important role in the macrophage clearance of nanoparticles. 
In humans, CR1 receptors are also found in monocytes and erythrocytes, 
while it is found on platelets in rats. Hence, these cells also play an important role 
in the clearance of immune complexes via the C3b-CR1 interaction. In humans, 
the erythrocyte-bound immune complexes are transferred to macrophages during 
their passage through the liver and the spleen; due to the multitude of erythrocytes, 
this may be a major contributor to the complement-mediated clearance 
process. 
In the case of liposomes, it has been demonstrated that the vesicle-blood protein 
interaction is largely determined by the fluidity and hydrophobicity of the vesicles, 
as well as by the vesicle size. Cholesterol free and cholesterol poor vesicles are 
rapidly captured by macrophages, mostly of the liver; the uptake is enhanced in 
the presence of serum. In contrast to hepatic phagocytes, bone marrow cells preferentially 
capture cholesterol-rich rather than cholesterol-poor vesicles primed with 
serum. On the other hand, serum stimulates the uptake of all cholesterol-containing 
vesicles by both splenic phagocytes and peritoneal macrophages, but its effect 
is significantly greater on the uptake of cholesterol-rich rather than cholesterolpoor 
vesicles. In addition, serum suppressed the uptake of vesicles prepared from 
Nanoparticulate Delivery to Reticuloendothelial System 467 
saturated phospholipids by liver phagocytes, but enhanced their uptake by spleen 
and bone-marrow macrophages. Thus, vesicle uptake by the liver, spleen, bonemarrow 
and peritoneal macrophages is determined by tissue-specific opsonins. 
Again, this is determined by the membrane fluidity and hydrophobicity which 
plays an important role in attracting the right opsonins which determine phagocytic 
activity. From the above observations, it has been suggested that a balance 
between the opsonins and dysopsonins (i.e. naturally occurring substances 
that inhibit phagocytic ingestion, usually by altering the surface properties of the 
phagocyte or particle or both, thereby interfering with opsonization or altering the 
metabolic activity of the phagocyte) may regulate the uptake of vesicles by phagocytes. 
Moreover, large vesicles (above 400 nm) are more readily cleared by liver 
macrophages probably by complement activation, while smaller ones (100-400 nm) 
localize in the spleen and bone marrow macrophages. Although, calcium at physiological 
levels is a prerequisite for the process of phagocytosis, it has been demonstrated 
that elevation of serum calcium levels above normal can inhibit, while a 
decrease below normal can facilitate the opsonophagocytosis of particles by Kupffer 
cells.5'6 
(b) Opsonin-independent recognition 
Non-opsonic blood proteins could also play an important role in particle clearance. 
Non-opsonic proteins, after adsorption onto particle surfaces, could experience conformational 
changes. Such changes are likely to expose chemical groups that could 
either be recognized directly by certain receptors on the phagocytic cell surface, 
or could act as ligands for the subsequent recognition by the blood opsonins. The 
molecules involved include mannose-binding protein (MBP), a C-type lectin with 
specificity for mannose, and N-acetyl glucosamine, also known to activate complement 
through the activation of Clr2 Cls2 complex, and by opsonization through 
macrophage Clq receptor. Scavenger receptors (SR) on macrophage plasma membranes 
and endothelial cells can recognize modified lipoproteins, polyanionic 
macromolecules, bacterial polysaccharides, silica and possibly anionic liposomes. 
FcyRII-B2 and FcyRI are regarded as putative receptors for low-density lipoproteins 
functioning, independent of IgG. CD14 is a physiologically important receptor for 
lipopolysaccharide, which is also a ligand for SR-AI. It is strongly expressed on 
monocytes and granulocytes, but on Kupffer cells, it is only expressed in chronic 
and acute liver diseases. A putative stearylamine receptor on Kupffer cells may 
also play a minor role in the clearance of neutral and anionic vesicles. Co-operation 
between different macrophage receptors, (e.g. fibronectin or immunoglobulins with 
complement or a y/33 with CD36) may increase the efficiency of particle phagocytosis 
and the clearance from blood. 
468 Basu & Lata 
Alternatively, macrophages as well as hepatocytes and liver endothelial cells 
may phagocytose/endocytose liposomes via direct recognition of phospholipid 
headgroups. Phospholipid recognition may be mediated by a number of plasma 
membrane receptors such as lectin receptors, CD14, various classes of scavenger 
receptors (e.g. classes A, B and D), FcyRI and FcyRII-B2.2 The recognition is specific 
for unsaturated phospholipids and fails for saturated phospholipids.7 
3.3. Factors influencing uptake 
Extrapolating the above discussion on liposomes to nanoparticles, it may be emphasized 
that some of the factors which affect their uptake by the cells of the RES 
include particle size, surface charge and surface hydrophobicity/hydrophilicity 
Cholesterol-free nanoparticles of diameter above 200 nm, incorporated with unsaturated 
phospholipid headgroups, are expected to be preferentially sequestered 
by liver macrophages and endothelial cells, while priming nanoparticles that 
are smaller than 200 nm diameter, containing phospholipid and probably cholesterol, 
with serum, may enhance their uptake by the spleen and bone marrow 
macrophages. The hepatic phagocytosis may be facilitated by subnormal blood 
calcium concentration. 
Although the above hypotheses have not yet been experimentally proven 
entirely, it has been demonstrated many times in our laboratory that poly- 
DL-lactide (PLA) and poly-DL-lactide co-glycolide (PLGA) nanoparticles of 
approximately 250 nm in diameter, containing a high percentage (58.8% w/w) of 
unsaturated phospholipids (phosphatidyl ethanolamine or phosphatidyl choline), 
are highly effective in reducing the spleen and liver parasite loads in the hamster 
models of experimental visceral leishmaniasis.8-10 This lends credence to the 
above view. 
In general, it appears that surface hydrophobic nanoparticles of size greater 
than 200 nm in diameter have a greater chance of being sequestered by macrophages 
of the liver and spleen. However, it has been reported that very small nanoparticles 
(< 70 nm diameter), consisting of poly-DL-lactide (PLA) — poly-ethylene-glycol 
(PEG) copolymeric micelles, can pass through the sinusoidal fenestrations in the 
liver and gain access to the liver parenchymal cells.11 Moreover, the effect of surface 
charge on phagocytosis and the biodistribution of albumin nanoparticles have 
been reported by Roser et al.n It has been noticed that albumin nanoparticles, with a 
zeta potential close to zero, showed a reduced in vitro phagocytic uptake by primary 
mouse peritoneal macrophages and a human hematopoietic cell line U-937, in comparison 
to charged particles, especially particles with a positive zeta potential. However, 
this difference has not been reflected in their in vivo blood circulation times 
and organ distributions in rats. Moreover, the influence of surface characteristics 
Nanoparticulate Delivery to Reticuloendothelial System 469 
which include surface charge density and zeta potential, along with size and surface 
hydrophobicity, has already been noticed on plasma protein adsorption patterns 
on colloidal drug carriers after intravenous administration, thus influencing their 
in vivo organ distribution.13 
3.4. Role of surface modifications on uptake 
Various surface modifications of nanoparticles have been shown to facilitate their 
uptake by different components of the RES. Colloidal gold nanoparticles, after 
opsonization with autologous plasma, are found to accumulate in Kupffer cells, 
the predominant opsonizing factor being fibronectin.14 Monocrystalline iron oxide 
nanoparticles (MION) were found to be readily captured by macrophages, and 
opsonization by the serum component C3, vitronectin and fibronectin resulted 
in a six-fold increase.15 Poly-lactide nanoparticles are sequestered by monocytes 
by passive adsorption and energy-requiring receptor-mediated endocytosis and 
the uptake are enhanced by opsonization with lipoproteins.16 Body distribution 
of fully biodegradable [14C]-poly (lactic acid) nanoparticles coated with albumin, 
after parenteral administration to rats, was examined by Bazile et al.17 As deduced 
from whole-body autoradiography and quantitative distribution experiments, the 
14C-labelled polymer is rapidly captured by the liver, bone marrow, lymph nodes, 
spleen and peritoneal macrophages. Nanoparticle degradation was addressed following 
14C excretion. The elimination of 14C was quick on the first day (i.e. 30% of 
administered dose), but slowed down subsequently. 
The block co-polymers of poloxamine and poloxamer series play an important 
role in the surface modification of nanoparticles. Of these, poloxamine 908 
is a poly-oxyethylene (POE)/polyoxypropylene (POP) block copolymer, which 
adsorbs on nanoparticles via the relatively hydrophobic POP segments, while the 
mobile POE chains extend outward, suppressing aggregation and providing stability. 
Poloxamer-407 is a block copolymer of POE /POP and a non-ionic surfactant. 
Polystyrene microspheres coated with the block copolymers of poloxamer 
and poloxamine series were observed to adsorb IgG, complement C3, transferrin 
and fibronectin in 50% serum, as well as fibrinogen in 50% plasma.18 Poloxamine 
908 activates the mononuclear phagocyte system so that the coated particles are 
sequestered by liver.19 Poly (organophosphazene) nanoparticles coated with poloxamine 
908 were mainly captured by the rat liver, while poly (organo phosphazene) 
nanoparticles coated with a novel poly (organophosphazene)-poly (ethyleneoxide) 
copolymer with a 5000 MW PEO chain were reported to be significantly 
targeted at the bone marrow.20 It has been observed that the sequestration of 
surface-engineered polystyrene nanospheres by the liver and spleen could be 
greatly augmented by the modification of poloxamer 407 and poloxamine 908, 
470 Basu & Lala 
by introducing a terminal protonated amine group to each PEO chain.21 Also, the 
phagocytic uptake of poloxamer and poloxamine coated polystyrene particles by 
mouse peritoneal macrophages was found to decrease with increasing adsorbed 
layer thickness, i.e. longer hydrophilic polymer chains of the coating agent, and 
consequently, a greater steric stabilization effect.22 
Surface engineered sterically stabilized nanospheres were synthesized and 
found to have enhanced drainage into lymphatics, as well as enhanced uptake by 
macrophages of the regional lymph nodes.23 Lymph node localization of biodegradable 
poly-lactide co-glycolide nanospheres could be enhanced by coating them 
with poloxamer and poloxamine.24 Correlation was observed between the length 
of the stabilizing POE chains of the block co-polymers polyoxyethylene (POE)/ 
polyoxypropylene (POP) in the poloxamer/poloxamine and nanosphere drainage 
and the passage across tissue-lymph interface in dermal lymphatic capillaries in 
the rat-footpads. The longer the POE chains, the faster the particle drainage. In 
order for effective opsonization of the nanospheres to occur in the lymphatics, the 
POE chains of the block copolymers should be of 5-15 ethylene oxide units. If 
the dimensions of the stabilizing POE chains of the poloxamines and poloxamers 
exceed the range of the Van der Waals forces of attraction, opsonization fails to occur 
and the surface modified nanospheres escape sequestration by the macrophages 
of the regional lymph nodes, and are rapidly drained into the systemic circulation, 
where they persist for prolonged periods.23 It has also been reported that 
polystyrene and poly-lactide-co-glycolide nanoparticles show enhanced localization 
in the lymph nodes when their surfaces are coated with polylactide (PLA) : 
poly-ethylene-glycol (PEG) or by producing co-precipitate nanospheres of PLGA 
and PLA : PEG, depending on surface characteristics.25 PEG-coated magnetite 
nanospheres have also been utilized to target diagnostic agents to regional lymph 
nodes.26 
It has been reported that small colloidal nanoparticles (< 150 nm in diameter) 
can be targeted specifically to the sinusoidal endothelial cells of the rabbit bone 
marrow, following intravenous administration, by coating their surface with the 
block co-polymer poloxamer 407, a non-ionic surfactant.27 
Influence of surfactant concentration on the body distribution of the nanoparticles 
was studied by Araujo et al.28 They noticed that the rapid RES uptake of the 
nanoparticles after intravenous injection, especially by the liver, can be reduced 
and the body distribution can be altered by coating them with non-ionic surfactants, 
e.g. poloxamine 908 and polysorbate 80. Evaluation of the likely mechanisms 
that contribute to the prolonged circulation times of sterically protected nanoparticles 
has already been made.29 Recent evidence showed that sterically stabilized 
particles are prone to opsonization, particularly by the opsonic components of the 
complement systems. 
Nanoparticulate Delivery to Reticuloendothelial System 471 
4. Active Targeting of Nanoparticles by 
Receptor Mediated Endocytosis 
Active targeting of nanoparticles to the organs of reticuloendothelial system could 
be done by attaching appropriate ligands for the well identified receptors on the 
target cells belonging to this system. Taking advantage of the presence of mannosyl 
fucosyl receptors on the macrophage surface, mannose bearing polymeric 
delivery systems have been designed and used with appropriate antileishmanial 
drug for site-specific delivery in the hamster model of experimental leishmaniasis.30 
These modified polymeric vesicles have been developed by coupling the amino 
group of phosphatidyl ethanolamine (PE), an essential compound of polymeric 
vesicles (PLGA) with amino group of p-aminophenyl a-D mannoside, in the presence 
of glutaraldehyde as a bridging agent (Fig. 1). The results demonstrate that 
because of receptor mediated endocytosis, nanoparticle entrapment of antileishmanial 
compound enhanced its effectiveness, an effect that seemed to be much 
greater when mannose bearing polymeric vesicles are used. Similarly, nanoparticles 
coated with the polymer mannan as ligand have been demonstrated to have a 50% 
enhanced uptake than uncoated nanoparticles by mannose-receptor positive mouse 
< & 
PE-vesicle 
NH2 + NH2 ^ Q | \ - Mannose 
p-aminophenyl a -D mannoside 
CHO-CH2-CH2-CH2-CHO 
glutaraldehyde 
N = CH-CH2-CH2-CH2—CH = N - { Q / - Mannose 
Mannose-grafted polymeric vesicles (a) 
NH2+ NH2—F(ab1)2 
PE-vesicle 
CHO- CH2- CH2- CH2-CHO 
glutaraldehyde 
N = CH-Cr-L-CH.— CH — CH = N — F (ab1)2 
Antibody - coated polymeric vesicles (b) 
Fig. 1. Formation of mannose-grafted polymeric vesicles (a) and antibody-coated 
mannose-grafted polymeric vesicles (b). 
472 Basu & Lata 
macrophage cell line (J774E), by the process of receptor mediated endocytosis.31 
Alternatively, the possibility of grafting a monoclonal antibody raised against a 
parasite-specific antigen onto the polymeric vesicle surface cannot be ignored for 
active targeting of an antileishmanial compound. Besides the grafting of the synthetic 
mannoside or the coating with the polymer mannan, similar results could be 
obtained when indigenous glycosides, e.g. Bacopasaponin C and Arjunglucoside I, 
both isolated from the Indian medicinal plants, Bacopa monniera and Terminalia bellerica 
respectively, having glucose at the terminal end of glycosidic chain (Fig. 2), 
are incorporated in PLA nanoparticles and are subjected to test for antileishmanial 
property, using both free and nanoparticle-incorporated forms.8,32 Much better 
therapeutic efficacy could be noticed with the polymeric vesicles incorporated with 
either of the two glycosides compared with the glycoside-unincorporated control 
vesicles. The unique presence of a glucose residue at the terminal end of the glycosidic 
chain, equipped the compounds to be self-targeting molecules that can be 
directed towards the glucose receptors present on the macrophage surface for facilitating 
a receptor mediated drug delivery to the target cells. Perhaps these are the 
Amarogenlin (MW 586) Bacopasaponin C (MW 898) 
Fig. 2. Structures of some glycosides isolated from indigenous sources. 
Nanoparticulate Delivery to Reticuloendothelial System 473 
very first reports for the targeted delivery of antileishmanial compounds in experimental 
leishmaniasis, a RES-associated disorder, using polymeric vesicles as drug 
carriers. 
5. Application in Chemotherapy 
Among the major RES associated disorders, pulmonary tuberculosis is identified 
as a killer disease because its death toll every year is enormously high. With a 
view to develop appropriate delivery systems to test the efficacy of frontline antitubercular 
drugs in vivo, experimental tuberculosis was induced in murine models 
and nanoparticle-encapsulated antitubercular drugs were administered orally to 
them at every 10th day When examined, no tubercle bacilli could be detected in 
the tissues after 5 such oral doses of treatment. Thus, nanoparticle encapsulated 
antitubercular drugs turned out to be a potential oral drug delivery system against 
murine tuberculosis.33 Alternatively, subcutaneous nanoparticle based antitubercular 
chemotherapy was also tried. Injectable PLG nanoparticles were found to 
hold promise for increasing drug bioavailability and reducing dosing frequency 
for a better management of tuberculosis.34 However, nebulization via the respiratory 
route of nanoparticle-based antitubercular drugs were reported to form a 
sound basis for improving drug bioavailability and reducing the dosing frequency 
for better chemotherapeutic control of pulmonary tuberculosis.35 
The next major RES associated disorder is leishmaniasis, which causes substantial 
human morbidity and mortality in many parts of the world. In an attempt 
to probe the disease, several new drugs as well as new delivery systems were put 
forward with a view to increase the drug efficacy and to reduce the drug toxicity. 
Using nanoparticle-bound pentamidine in a Leishmania major /mouse model, ultra 
structural changes in parasites were noticed by Fusai et al.36 In the parasites inside 
the Kupffer cells, transmission electron microscopy showed a swollen mitochondrion 
with a loss of cristae, destruction or fragmentation of the kinetoplast, loss 
of ribosomes and the destruction of parasite structures except for the subpellicular 
microtubules. The therapeutic efficacy of several indigenous antileishmanial 
agents e.g. Bacopasaponin C, isolated from the Indian medicinal plant Bacopa monniera, 
Quercetin, isolated from Fagopyrum esculentum and Harmine, isolated from 
Peganum harmala, were not only studied but compared after incorporating them in 
different vesicular delivery systems against experimental leishmaniasis in hamster 
models.8-10 At equivalent quercetin9 concentration, the nanocapsulated quercetin 
was found to be the most potent in reducing the parasite burden in the spleen as 
well as in reducing hepatotoxicity and renal toxicity, compared with free drug or 
drug in other vesicular forms. Similarly, Bacopasaponin C8, at an equivalent dose of 
1.75 mg/kg body weight and Harmine,10 at an equivalent dose of 1.5 mg/kg body 
474 Basu & Lata 
weight, were found to be very active in all the vesicular forms, but the best efficacy 
in the lowering of spleen parasite load was found with the nanocapsulated form. 
Thus, in each case, the nanoparticle-loaded antileishmanial agent was found to be 
most efficient in the lowering of spleen parasite load and the efficacy was found to 
vary in the following order: 
Nanoparticles > niosomes > liposomes > microspheres > free drug 
and the hepatotoxicity, as well as the renal toxicity was found to follow in the reverse 
order as shown above. In vitro antileishmanial activity of amphotericin B loaded in 
poly (epsilon-caprolactone) nanospheres was also noted, but the nanospheres did 
not show any improvement of amphotericin B activity against the resistant strain.37 
Attempt was made to deliver piperine to treat experimental visceral leishmaniasis 
in mice model using oil in water emulsions known as lipid nanospheres (LN) 
or fat emulsions. A single dose of 5 mg/kg of lipid nanospheres of piperine was 
found to significantly reduce the liver and splenic parasite burden.38 Therapeutic 
evaluation of free and nanocapsule encapsulated atovaquone was made in the 
treatment of murine visceral leishmaniasis by Cauchetier et al.39 The liver parasite 
burdens, evaluated by using the Stauber method, indicated that the atovaquoneloaded 
nanocapsules were significantly more effective than the free drug. 
Trypanosomiasis, another deadly disease caused by the parasite Trypanosoma 
cruzi was also challenged by using nanoparticles of polyalkylcyanoacrylate as a 
targeted delivery system for nifurtimox. The drug-loaded nanoparticles significantly 
increased trypanocidal activity compared with the empty one.40 The use 
of poly (lactic-coglycolic acid) nanoparticles for targeted oral drug delivery to the 
inflamed gut tissue in the inflammatory bowel disease was examined.41 Such a 
strategy of local drug delivery was considered to be a distinct improvement, compared 
with existing colon delivery devices for this disease. Efficacy of nanoparticles 
as carrier systems for antiviral agents in human immunodeficiency virus-infected 
human monocytes/macrophages was evaluated in vitro by Bender et al.42 In the 
same year, macrophage targeting of antivirals, e.g. azidothymidine, was evaluated 
in vivo as a promising strategy for AIDS therapy.43 The authors, after analyzing 
the results, concluded that nanoparticles could be considered as a promising drug 
targeting system for azidothymidine to the RES organs. They also hypothesized 
that an increase in drug availability at the sites containing abundant macrophages 
might allow a reduction in dosage in order to avoid systemic toxicity. 
For targeted gene delivery, calcium phosphate nanoparticles were found to be 
a unique class of non-viral vectors, which can serve as efficient and alternative 
DNA carriers. Moreover, the surface of these nanoparticles was suitably modified 
by absorbing a highly adhesive polymer e.g. polyacrylic acid and these surface 
Nanoparticulate Delivery to Reticuloendothelial System 475 
modified calcium phosphate nanoparticles were used in vivo to target genes specifically 
to the liver.44 Chitosan-DNA nanoparticles were designed as gene carriers 
using a complex coacervation process. The transfection efficiency was found to be 
cell-type dependent.45 Conjugation of PEG on the nanoparticles allowed lyophilization 
without aggregation and without the loss of bioactivity. The clearance in mice 
following intravenous administration was found to be slower than unmodified 
nanoparticles, with a higher deposition in kidney and liver. Use of sodium chloride 
modified silica nanoparticles (SNAP) as a novel non-viral vector with a high efficiency 
of DNA transfer into cells has already been reported.29 Previous gene transfer 
methods using non-viral vectors, such as liposomes or nanoparticles, resulted in 
relatively low levels (35 to 50% approx.) of gene expression. SNAP showed a better 
efficiency (about 70%) of DNA transfection into cells, as well as a better protection 
of DNA against degradation. Intravenous and/or intra-abdominal administration 
of the SNAP to mice revealed the accumulation of SNAP in the cells of the brain, 
liver, spleen, lung, kidney, prostate and testis, without any pathological cell changes 
or mortality, suggesting that they passed through the blood-brain, blood prostate 
and blood-testis barriers. 
Sponge-like alginate nanoparticles were found to be a new potential system for 
the delivery of antisense oligonucleotides. The aim of this study was to design a new 
antisense oligonucleotide (ON) carrier system based on alginate nanoparticles, and 
to investigate its ability to protect ON from degradation in the presence of serum. 
From the results, such nanosponges were found to be promising carriers for specific 
delivery of ON to the lung, liver and spleen.46 
6. Summary 
During the last few decades, numerous approaches have been explored to modify 
the biodistribution and bio-availability of the drugs by using carriers systems of 
colloidal dimensions, e.g. liposomes, micelles and nanoparticles. From the various 
studies reported in the literature, it can be concluded that the factors responsible for 
particle uptake are the particle size, their surface charge, surface hydrophobicity 
and the presence and/or absence of surface ligands. Keeping these key factors in 
mind, the designing and production of polymeric nanoparticles has been investigated 
since the late 1970s. The major challenge in the development of particulate 
carriers for targeting at specific body sites is the preparation of the particles of optimum 
size with hydrophilic surfaces so as to have long circulation time in the blood 
and escape from RES scavenging. The body's RES, mainly the Kupffer cells in the 
liver usually take up polymeric nanoparticles with hydrophobic surface. Therefore, 
the residence time of these nanoparticles in the blood is considerably small. 
However, as it has been observed that nanoparticles such as other colloidal drug 
476 Basu & Lata 
delivery systems, on intravenous injection, are rapidly sequestered and retained by 
the organs comprising the reticuloendothelial system (RES), so that the targeting 
of nanoparticles to RES is a much simpler process than the targeting to any other 
organ. 
The major defense system of the body, i.e. the reticuloendothelial system or 
more correctly, the mononuclear phagocyte system can recognize any foreign elements 
(here the injected nanoparticles) through the opsonization process. The 
Kupffer cells (macrophages) of the liver and of course to a lesser extent, the 
macrophages of the spleen and the circulating macrophages play an important 
role in removing the opsonized particles. Particle size and surface properties of 
the particles can modulate the process of particle capture. Particles that have large 
hydrophobic surface are efficiently coated with plasma components and are rapidly 
removed from circulation. Thus, injected nanoparticles are covered by plasma proteins 
immediately. The larger particles are trapped in the liver but the smaller ones 
can reach the general circulation and the modified surfaces can be directed to the 
inflammation sites, endothelial cells or spleen. Targeting, usually achieved by injecting 
nanoparticles in vivo, is mainly passive, although active targeting is being done 
very recently. An excellent example of passive targeting is the uptake of nanoparticles 
by the Kupffer cells of liver. In many cases, this targeting can be exploited 
to help treatment in disease conditions e.g. leishmaniasis and candiasis, where 
macrophages are directly involved in the disease processes. For greater specificity, 
the active targeting of the nanoparticulate delivery systems can be achieved by 
attaching the targeting ligand, appropriate to the receptors on the target cells, to 
the surface of the particle conjugate. Monoclonal antibodies and sugar residues 
are the possible ligands. The hepatocytes in the liver is an important target site 
for some diseases such as hepatitis, as well as in gene therapy. In gene therapy, 
the liver can be used as "bioreactors" where the administered gene can be used to 
express the missing factors such as growth hormones and blood factors. In the liver, 
the endothelial lining of the blood vessels (sinusoids) have gaps or fenestrations, 
through which nanoparticle can pass and there is no intact basement membranes 
below these fenestrations. Thus, the nanoparticles can have a close interaction with 
the liver hepotocytes. The size of the gap was estimated to be between 100 nm and 
150 nm. Hence, recent work in the field has suggested that the size of the nanoparticles 
should be less than 50 ran in diameter for better interaction with the hepatocytes. 
The polymeric nanoparticles, besides being biocompatible and biodegradable 
and having longer circulation time in blood, remain unaffected by circulating 
lipases that protect the drug from the bioenvironment. In an attempt to acquaint the 
readers with the sequence of events that are associated with nanoparticulate drug 
delivery to the RES-associated disorders, this chapter first identifies the reticuloendothelial 
systems (RES), discusses about the possible mechanisms of the uptake of 
Nanoparticulate Delivery to Reticuloendothelial System 477 
nanoparticles by them, and finally, updates the application of drug-loaded nanoparticles 
in the chemotherapy of diseases associated with RES. Moreover, this chapter 
contributes to the furtherance of our present knowledge in the area of targeting by 
suggesting that the composition, surface characteristics and the size of the delivery 
vesicles are the three important parameters that must be considered when drawing 
a strategy for efficient delivery. 
Acknowledgment 
The authors gratefully acknowledge the financial support provided to M.K.Basu 
by the Council of Scientific and Industrial Research (CSIR), Government of India, 
in the form of the Emeritus Scientist scheme. 
References 
1. Moghimi SM and Patel HM (1998) Serum mediated recognition of liposomes by phagocytic 
cells of the reticuloendothelial system — The concept of tissue specificity. Adv Drug 
Del Rev 32:45-60. 
2. Moghimi SM and Hunter AC (2001) Recognition by macrophages and liver cells of 
opsonized phospholipid vesicles and phospholipid head groups. Pharm Res 18(l):l-8. 
3. Absolom D (1986) Opsonins and dysopsonins — An overview. Meth Enzymol 
132:323-326. 
4. Hostetter MK, Krueger RA and Schmelling DJ (1984) The biochemistry of opsonization : 
Central role of the reactive thiolester of the third component of complement. / Infect Dis 
150:653-661. 
5. Ryder KW Jr, Kaplan JE and Saba TM (1975) Serum calcium and hepatic Kupffer cell 
phagocytosis. Proc Soc Exp Biol Med 149:163-167. 
6. Moghimi SM and Patel HM (1990) Calcium as a possible modulator of Kupffer cell 
phagocytic function by regulating liver specific opsonic activity. Biochim Biophys Acta 
1028(3):304-308. 
7. Moghimi SM and Patel HM (1989) Serum opsonins and phagocytosis of saturated and 
unsaturated phospholipid liposomes. Biochim Biophys Acta 984(3):384-387. 
8. Sinha J, Raay B, Das N, Medda S, Garai S, Mahato SB and Basu MK (2002) Bacopasaponin 
C: Critical evaluation of antileishmanial properties in various delivery modes. Drug Del 
9:55-62. 
9. Sarkar S, Mandal S, Sinha J, Mukhopadhyay S, Das N and Basu MK (2003) Quercetin : 
Critical evaluation as an antileishmanial agent in vivo. } Drug Targ 10:573-578. 
10. Lala S, Pramanick S, Mukhopadhyay S, Bandyopadhyay S and Basu MK (2004) Harmine: 
Evaluation of its antileishmanial properties in various delivery systems. / Drug Targ 
12(3):165-175. 
11. Stolnik S, Heald CR, Neal J, Garnett MC, Davis SS, Ilium L, Purkis SC, Barlow RJ and 
Gellert PR (2001) Poly-lactide poly(ethylene glycol) miceller-like particles as potential 
478 Basu & Lata 
drug carriers: Production, colloidal properties and biological performance. / Drug Targ 
95(5):361-378. 
12. Roser M, Fischer D and Kissel T (1998) Surface- modified biodegradable albumin nanoand 
microspheres II. Effect of surface charges on in vitro phagocytosis and biodistribution 
in rats. Eur J Pharm Biopharm 46(3):255-263. 
13. Luck M, Paulke BR, Schroder W, Blunk T and MuUer RH (1991) Analysis of plasma 
protein adsorption on polymeric nanoparticles with different surface characteristics. 
/ Biomed Mater Res 39(3):478-485. 
14. Moghimi SM, Porter CJH, Muir IS, Ilium L and Davis SS (1991) Non-phagocytic uptake 
of intravenously injected microspheres in rat spleen: Influence of particle size and 
hydrophilic coating. Biochem Biophys Res Commun 177:861-866. 
15. Moore A, Weissleder R and Bogdanov A Jr (1997) Uptake of dextran-coated monocrystalline 
iron oxides in tumor cells and macrophages. JMagn Reson Imaging 7(6):1140-1145. 
16. Leroux JC, Gravel P, Balant L, Volet B, Anner BM, Allemann E, Doelker E and Gurny 
R (1994) Internalization of poly-(D,L-lactic acid) nanoparticles by isolated human leucocytes 
and analysis of plasma proteins adsorbed onto the particles. / Biomed Mater Res 
28(4):471^81. 
17. Bazile DV, Ropert C, Huve P, Verrecchia T, Marland M, Frydman A, Veillard M and 
Spenlehauer G (1992) Body distribution of fully biodegradable [14C]-poly (lactic acid) 
nanoparticles coated with albumin after parenteral administration to rats. Biomaterials 
13(15):1093-1102. 
18. Norman ME, Williams P and Ilium L (1993) Influence of block copolymers on the adsorption 
of plasma proteins to microspheres. Biomaterials 14(3):193-202. 
19. Armstrong TI, Moghimi SM, Davis SS and Ilium L (1997) Activation of the mononuclear 
phagocyte system by poloxamine 908: Its implication for targeted drug delivery. Pharm 
Res 14(11):1629-1633. 
20. Vandorpe J, Schact E, Dunn S, Hawley A, Stolnik S, Davis SS, Garnett MC, Davies MC 
and Ilium L (1997) Long-circulating biodegradable poly-(phosphazene) nanoparticles 
surface-modified with poly-(phosphazene)-poly(ethylene oxide) copolymer. Biomaterials 
18(17):1147-1152. 
21. Neal JC, Stolnik S, Garnett MC, Davis SS and Ilium L (1998) Modification of the copolymers 
poloxamer 407 and poloxamine 908 can effect the physical and biological properties 
of surface-modified nanospheres. Pharm Res 15(2):318-324. 
22. Ilium L, Jacobson LO, Muller LH, Mok E and Davis SS (1987) Surface characteristics and 
the interaction of colloidal particles with mouse peritoneal macrophages. Biomaterials 
8(2):113-117. 
23. Moghimi SM, Hawley AE, Christy NM, Gray T, Ilium L and Davis SS (1994) Surface 
engineered nanospheres with enhanced drainage into lymphatics and uptake by 
macrophages of the regional lymph nodes. FEBS Lett 344(l):25-30. 
24. Hawley AE, Ilium L and Davis SS (1997) Lymph node localization of biodegradable 
nanospheres surface modified with poloxamer and poloxamine block copolymers. FEBS 
Lett 400(3):319-323. 
Nanoparticulate Delivery to Reticuloendothelial System 479 
25. Hawley AE, Ilium L and Davis SS (1997) Preparation of biodegeradable, surfaceengineered 
PLGA nanospheres with enhanced lymphatic drainage and lymph node 
uptake. Pharm Res 14(5):657-661. 
26. Ilium L, Church AE, Butterworth MD, Arien A, Whetstone J and Davis SS (2001) Development 
of systems for targeting the regional lymph nodes for diagnostic imaging: In vivo 
behavior of colloidal PEG-coated magnetite nanospheres in the rat following interstitial 
administration. Pharm Res 18(5):640-645. 
27. Porter CJ, Moghimi SM, Ilium L and Davis SS (1992) The poly oxyethylene/
polyoxypropylene block co-polymer poloxamer-407 selectively redirects intravenously 
injected microspheres to sinusoidal endothelial cells of rabbit bone-marrow. 
FEBS Lett 305(l):62-66. 
28. Araujo L, Lodenberg R and Kreuter J (1999) Influence of the surfactant concentration on 
the body distribution of nanoparticles. / Drug Targ 6:373-385. 
29. Chen Y, Xou Z, Zheng D, Xia K, Zhao Y, Liu T, Long Z and Xia J (2003) Sodium chloride 
modified silica nanoparticles as a non-viral vector with a high efficiency of DNA transfer 
into cells. Curr Gene Ther 3:273-279. 
30. Medda S, Jaisankar P, Manna RK, Pal B, Giri VS and Basu MK (2003) Phospholipid 
microspheres: A novel delivery mode for targeting antileishmanial agent in experimental 
leishmaniasis. / Drug Targ 11(2): 123-128. 
31. Cui Z, Hsu CH and Mumper RJ (2003) Physical characterization and macrophage cell 
uptake of mannan-coated nanoparticles. Drug Dev Ind Phar 29(6):689-700. 
32. Tyagi R, Lala S, Verma AK, Nandy AK, Mahato SB, Maitra A and Basu MK (2005) Targeted 
delivery of arjunglucoside I using surface hydrophilic and hydrophobic nanocarriers to 
combat experimental leishmaniasis. / Drug Targ 13:161-171. 
33. Pandey R, Zahoor A, Sharma S and Khuller GK (2003) Nanoparticle encapsulated antitubercular 
drugs as a potential oral drug delivery system against murine tuberculosis. 
Tuberculosis 83:373-378. 
34. Pandey R and Khuller GK (2004) Subcutaneous nanoparticle-based antitubercular 
chemotherapy in an experimental model. / Antimicrob Chemother 54:266-268. 
35. Pandey R, Sharma A, Zahoor A, Sharma S, Khuller GK and Prasad B (2003) Poly (DLlactide-
co-glycolide) nanoparticle-based inhalable sustained drug delivery system for 
experimental tuberculosis. / Antimicrob Chemother 52:981-986. 
36. Fusai T, Boulard Y, Durand R, Paul M, Bories C, Rivollet D, Astier A, Houin R and 
Deniau M (1997) Ultrastructural changes in parasites induced by nanoparticle-bound 
pentamidine in a Leishmania major/'mouse model. Parasite 4:133-139. 
37. Espuelus MS, Legrand P, Loiseau PM, Bories C, Barratt G and Irache JM (2002) 
In vitro antileishmanial activity of amphotericin B loaded in poly (epsilon-caprolactone) 
nanospheres. / Drug Targ 10:593-599. 
38. Veerareddy PR, Vobalaboina V and Nahid A (2004) Formulation and evaluation of oilin-
water emulsions of piperine in visceral leishmaniasis. Pharmazie 59:194-197. 
39. Cauchetier E, Paul M, Rivollet D, Fessi H, Astier A and Deniau M (2003) Therapeutic 
evaluation of free and nanocapsule-encapsulated atovaquone in the treatment of murine 
visceral leishmaniasis. Ann Trop Med Parasitol 97:259-268. 
480 Basu & Lata 
40. Gonzalez-Martin G, Merino I, Rodriguez-Cabezas MN, Torres M, Nunez R and Osuna A 
(1998) Characterization and trypanocidal activity of nifurtimox-containing and empty 
nanoparticles of polyethyl cyanoacrylates. / Pharm Pharmacol 50:29-35. 
41. Lamprecht A, Ubrich N, Yamamoto H, Schafer U, Takeuchi H, Maincent P, Kawashima Y 
and Lehr CM (2001) Biodegradable nanoparticles for targeted drug delivery in treatment 
of inflammatory bowel disease. / Pharmacol Exp Ther 299:775-781. 
42. Bender AR, Von Briesen H, Kreuter J, Duncan IB and Rubsamen-Waigmani H 
(1996) Efficiency of nanoparticles as a carrier system for antiviral agents in human 
immunodeficiency-virus infected human monocytes /macrophages in vitro. Antimicrob 
Agents Chemother 40:1467-1471. 
43. Lodenberg R and Kreuter J (1996) Macrophage targeting of azidothymidine: A promising 
strategy for AIDS therapy. AIDs Res Hum Retroviruses 12:1709-1715. 
44. Roy I, Mitra A, Moitra A and Mozumdar S (2003) Calcium phosphate nanoparticles as 
novel non-viral vectors for targeted gene delivery. Int} Pharm 250:25-33. 
45. Mao HQ, Roy K, Troung-Le VL, Janes KA, Lin KY, Wang Y, August JT and Leong KW 
(2001) Chitosan-DNA nanoparticles as gene carriers: Synthesis, characterization and 
transfection efficiency. / Control Rel 70(3):399-421. 
46. Aynie I, Vauthier C, Chacun H, Fattal E and Couvreur P (1999) Spongelike alginate 
nanoparticles as a new potential system for the delivery of antisense oligonucleotides. 
Antisense Nucleic Acid Drug Del 9:301-312. 
22 
Delivery of Nanoparticles to the 
Cardiovascular System 
Ban-An Khaw 
1. Introduction 
Nanoparticles have become one of the highly desirable drug delivery vehicles in 
recent years,1 not only due to the capacity but also due to their longevity. Most 
nanoparticles in use today are solid nanoparticles. Their applications in biological 
systems have both advantages as well as adverse effects.2 However, biocompatible 
nanoparticles such as liposomes or micells have circumvented some of these 
adverse consequences.1 Hood et al. reported the use of lipid based nanoparticles 
(40-50 nm), targeted with organic av/J3 ligands, to target the endothelium of tumor 
vasculature to induce anti-angiogenesis, following the delivery of mutant Raf gene.3 
In 2003, Kralj and Pavelic wrote, "the main interest currently lie in improving diagnostic 
methods and in developing better drug delivery systems to improve disease 
therapy" relative to the application of nanotechnology.4 The current chapter will be 
restricted to review of the application of nanoparticles, primarily nano-lipid vesicles 
subsequently referred to by the original name, liposomes, to the cardiovascular 
system, from diagnostic to therapeutic applications including novel cell membrane 
lesion sealing to gene delivery. 
2. Targeting the Myocardium with Immunoliposomes 
The interest in the use of nanoparticles, such as liposomes, for targeting the cardiovascular 
system has increased dramatically in recent years. The first application 
481 
482 Khaw 
of non-target specific liposomes for localization in experimental myocardial infarction 
was reported by Caride et al.5 They showed that plain cationic liposomes 
localized in the infarct better than neutral or anionic liposomes. However, the first 
targeted delivery of liposomes in cardiovascular application was reported by us 
in 1979.6 Although the exact size of the immunoliposomes used in that study was 
not determined, both multilamellar and unilamellar immunoliposomes were generated. 
These liposomes were target-specific and were demonstrated to be able to 
target cardiac myosin, exposed to the extracellular milieu following experimental 
acute myocardial infarction. Using the canine model, In-Ill labeled antimyosin 
immunoliposomes were demonstrated to localize in the infarct by gamma scintigraphic 
imaging, after catheter infusion of the immunoliposomes into the infarcted 
region. This study demonstrated the first potential application of liposomes as targeted 
nano-lipid vesicles for the delivery of various pharmaceuticals. 
Despite this potential for in vivo targeted drug delivery, it was observed that 
these immunoliposomes also had high non-target organ activities in vivo. Organs 
such as the liver and bone with high reticuloendothelial distribution were prime 
non-target organs for non-specific immunoliposomes sequestration. Therefore, we 
reasoned that if the antimyosin immunoliposomes were made to mimic normal cells 
such as lymphocytes, then these modified immunoliposomes might circumvent the 
affinity for the reticuloendothelial system. To mimic normal cells, sialoglycoprotein, 
fetuin, was attached to liposomes by either glutaraldehyde cross-linkage or cholate 
dialysis method in the presence or absence of immunoglobulins.7 Although the initial 
studies were tantalizing, unequivocal demonstration of this phenomenon was 
not achieved. The only clear cut advantage of sialoglycoprotein modified liposomes 
over plain liposome in mice was the increase retention of the liposome in the blood 
of mice at 15 min post intravenous administration (54.7 ± 11.0 vs 41.8 ±5.2% injected 
dose per gram, respectively). 
Subsequently, Klibanov and co-workers8 developed a method to prolong 
in vivo circulation time of liposomes, by coating liposomes with polyethyleneglycol. 
Torchilin et al. applied this method of polyethyleneglycol protection from 
sequestration by the reticuloendothelial system on antimyosin immunoliposomes, 
and demonstrated that PEG-antimyosin-immunoliposome with 10% mole PEG 
had slower blood clearance than PEG-antimyosin immunoliposomes with 4% mol 
PEG or just immunoliposomes.9 The half life (T1/2) of immunoliposomes in rabbits 
with experimental acute myocardial infarction was 40 min, whereas the T1/2 
of PEG-coated immunoliposomes at 10% mole PEG was about 1000 min (16hrs 
40 min) and 4% mol PEG was 200 min. This increase in circulation time enabled 
enhanced targeting of radiolabeled PEG-immunoliposomes in acute myocardial 
infarcts. The maximum ratio of infarct to normal tissue for plain liposomes was 
about 4:1, whereas that of 4% mol PEG-immunoliposomes was 20:1 and that of 
Delivery of Nanoparticles to the Cardiovascular System 483 
10% mole PEG-immunoliposomes was 12:1 at 6hrs post intravenous liposome or 
immunoliposomes delivery. The reduction in the uptake ratio of 10% mole-PEGimmunoliposomes 
is consistent with higher blood activity at the time of sacrifice 
than with 4% mole PEG-immunoliposomes. If the experiments were carried on 
longer, absolute uptake in the infarct, as well as the ratios of infarcted tissue to normal 
with 10% mole PEG-immunoliposomes should become greater than the values 
for the 4% mole PEG-immunoliposomes. 
Torchilin et al. later reported that size also affected the targeting potential 
of PEG-modified immunoliposomes in rabbits with experimental myocardial 
infarction.10 It appeared that small PEG-modified-antimyosin immunoliposomes 
of about 135 (120-150)nm diameter size had the highest accumulation of the 
intravenously administered immunoliposomes in the target (0.25 ± 0.14% injected 
dose per gram of tissue ± SD). Non-specific uptake of the same PEG-antimyosin 
immunoliposomes in normal myocardium was only 0.02 ±0.1% ID/g. Unmodified 
plain liposomes and PEG-modified plain-liposomes had 0.02 ± 0.01 and 
0.13 ± 0.10% ID/g respectively in the infarcted myocardium. Normal myocardial 
activity was respectively 0.004 ± 0.001 and 0.017 ± 0.006. Antimyosin-liposome 
injection resulted in 0.14 ±0.05 and 0.007 ±0.002% ID/g localization in the 
infarct and normal myocardium respectively. This resulted in the target to normal 
myocardial activity ratios of 5.17 ± 2.35 for plain liposomes, 8.05 ± 5.03 for 
PEG-plain liposomes, 22.70 ± 2.38 for antimyosin liposomes and 14.10 ± 7.15 for 
PEG-antimyosin immunoliposomes. The lower target to non-target ratio of PEGantimyosin 
immunoliposomes, relative to anti-myosin-immunoliposomes in the 
infarct, is due to the higher blood activity of the former at 5hrs post intravenous 
administration of liposome preparations (0.35 ±0.11 vs 0.060.01% ID/g 
respectively). 
When larger liposome preparations (350-400 nm diameter) were used, it was 
observed that plain liposomes had the same infarct localization activity (0.02 ± 0.01) 
as the small plain liposomes. However, modification with antimyosin, or with both 
antimyosin and PEG, resulted in lower target activity (0.09 + 0.04, 0.0.15 + 0.02 
respectively) but similar background activity (0.003 + 0.001 and 0.02 + 0.004 
respectively). It was reasoned that the lower targeting potential with the larger 
immunoliposomes was due to the limitation of these larger nanoparticles to 
extravasate into the extracellular interstitial matrices, even though the blood activities 
at 5 hrs were similar (larger PEG-immunoliposomes = 0.41 + 0.08 and small 
PEG immunoliposomes = 0.35 ± 0.11). The larger PEG-modified plain liposomes 
appeared to have similar non-target organ activity as the smaller PEG-liposomes. 
The mechanism of non-specific accumulation with PEG-modified plain-liposomes 
may be related to blood activity that allowed longer contact with non-target tissues 
when PEG modified large and small liposomes were used (0.38 + 0.02 and 
484 Khaw 
0.50 + 0.11 respectively). Both large and small liposomes had Ti/2s of 10 to 15 min. 
When they were modified with antimyosin, the Ti/2s were also similar between 
15-20 min. Small PEG-modified liposomes had a T1/2 of > 1000 min, whereas 
larger PEG-modified liposomes or small and large PEG-immunoliposomes had 
T1/2S > 600 min. It appears that this increase in blood circulation activity raised the 
background activity, as well as the absolute target activity, when PEG-modified 
antimyosin immunoliposomes were used. It was concluded that for diagnostic 
applications where high target to non-target activity is desirable, immunoliposomes 
would be the best candidate for use. However, in therapeutic applications 
where absolute concentration of the targeting agent determines the efficacy of the 
intervention, small PEG-immunoliposomes would be preferable. However, large 
PEG-immunoliposomes may also be useful due to the larger pay-load capacity of 
the larger liposomes, despite their lower target activity. 
3. Other Nanoparticle-Targeting of the 
Cardiovascular System 
Although nanoparticles have been used as targeting agents for tumors, blood and 
lymphatic vessels, the ultimate utility of such agents in the cardiovascular system 
is just beginning, even though the first in vivo demonstration of the feasibility 
of immunoliposome-application in the cardiovascular system was reported 
in 1979.6 Recently, Lanza and colleagues11 reported targeting of antiproliferative 
drugs, such as doxorubicin and paclitaxel, to the vascular smooth muscle cells 
in cell cultures with a magnetic resonance imaging nanoparticle contrast agent. 
In this study, the investigators prepared perfluorcarbon nanoparticles containing 
gadolinium-DTPA-bis-oleate in 2% surfactant comixture of lecithin and cholesterol. 
The resultant nanoparticles had a mean diameter of 250 nm. These nanoparticles 
were targeted using a three step procedure.12 Initial targeting was achieved with 
biotinylated monoclonal antibody to tissue factor (TF), followed by administration 
of avidin that bound to biotin. The third step consisted of administration of 
biotinylated non-gaseous, lipid-encapsulated, perflurocarbon emulsion nanoparticles 
loaded with doxorubicin or paclitaxel. The study showed that TF-targeted 
doxorubicin or paclitaxel loaded nanoparticles were more efficient antiproliferative 
agents than control targeted or non-targeted nanoparticles without drug loading. 
The same group also showed that in vivo targeting with antifibrin antibody 
enabled visualization of the fibrin clots in canine femoral arteries by intravascular 
ultrasound imaging.12 
Another application of targeted nanoparticles in the cardiovascular system was 
reported by Spragg et alP Their study showed that immunoliposomes sporting 
monoclonal antibody specific for an extracellular domain of E-selectin targeted 
Delivery of Nanoparticles to the Cardiovascular System 485 
human umbilical vein endothelial cells (HUVEC) only after activation of these cells 
with recombinant human interleukin 1/3. Localization of the targeted immunoliposomes 
was 13 to 275 fold higher in IL-l/J activated HUVEC than in unactivated ones. 
Other investigators have also shown that targeting of other adhesion molecules, 
such as ICAM-1, with antibodies to ICAM-1 was feasible in vitro.u Echogenic 
immunoliposomes targeted with antibodies to ICAM-1, VCAM-1, fibrin and tissue 
factors have recently been reported for imaging of atheroma in Yucatan miniswine 
model of endothelial denudation by intravascular ultrasound imaging.15 
Expression of ICAM-1, VCAM-1 and tissue factor, as well as fibrin deposition, 
were visualized within 5 min of antibody-targeted echogenic immunoliposomes 
administration. 
4. Novel Application of Nano-lmmunoliposomes 
Although most of the applications of nanoparticle size immunoliposomes in the 
cardiovascular system have been in imaging or targeted drug delivery, in 1995, we 
reported a novel application of antimyosin immunoliposomes for cell membrane 
lesion sealing of hypoxic cardiocytes.16 In this application, we reasoned that cell 
membrane lesions that develop in myocardial injury and ischemia in vivo or hypoxia 
in vitro result in irreversible myocyte death. However, if these cell membrane lesions 
were sealed prior to extensive loss of intracellular contents, and hypoxia or ischemia 
is removed, then the injured cells, with the lesions now sealed with a membrane 
sealing agent, should be able to remain viable and undergo membrane repair. This 
hypothesis is demonstrated in Fig. 1. The agent of cell membrane lesion sealing was 
proposed to be antimyosin immunoliposomes.16 The concept of cell membrane 
lesion sealing as a repair mechanism is not exclusive to our hypothesis. Many 
Anctionmj (.ML (« friqwiPii myosin 
tiirougti tnembfaiip li>SHm 
Fig. 1. Diagrammatic representation of the process of cell membrane lesion sealing with 
antimyosin immunoliposomes (CSIL). 
486 Khaw 
cells, including mammalian cells, undergo rapid self-sealing of the ruptured cell 
membrane.17-21 This is an innate property of many cells that responds to exposure 
to higher physiological concentration of Ca++ in the extracytoplasmic environment 
when lesions develop in the cell membrane, utilizing intracellular membrane 
vesicles such as lysosomes and endosomes to seal the lesions. This innate mechanism, 
although highly useful, is not sufficient when development of cell membrane 
lesions is more extensive. 
In our initial report, embryonic cardiocytes in tissue culture were used to 
demonstrate the role of antimyosin-immunoliposomes as Cytoskeletal-antigen 
Specific ImmunoLiposomes (CSILs) for sealing of cell membrane lesions induced 
by vigorous process of induction of hypoxia.16 Cells (2 x 106) in sterile 25 ml culture 
flasks with or without CSILs were flushed with N2 gas for 4 min vigorously 
into the media dislodging the cells. The caps were closed tight and the flasks were 
incubated in a 37°C 5% CO2 incubator for 24hrs. The viability of the cells were 
either assessed by trypan blue exclusion method or by [3H]thymidine uptake, after 
an additional 24 hrs of normoxic culture of the experimental cultures.16 Figure 2 
(left and right) shows that the viability of hypoxic cells treated with immunoliposomes 
(CSILs) (96.17 ± 1.24% by trypan blue exclusion or 3.26 + 0.483 x 106 cpm by 
[3H]thymidine uptake) was not significantly different from the viability of normoxic 
cultured controls (98.3 ± 0.58% or 3.68 ± 0.328 x 106 cpm respectively), whether 
viability was assessed by the dye exclusion or [3H]thymidine uptake method. Viability 
of the CSIL treated cells was significantly greater than the viability of hypoxic 
embryonic cardiocytes treated with plain liposomes (PL-Hypoxia, 42.3 ± 3.11% or 
1.14 ± 0.577 x 106 cpm), IgG-liposomes (IgGL-Hypoxia, 42.85 ± 6.24%), or hypoxia 
alone (13.97 ± 1.77% or 0.115 ± 0.155 x 106) (Fig. 2 left panel). Viability of controls 
Viability 
total 
0 0 - 
80 — 
6 0 - 
4 0 - 
2 0 - 
0 — 
Trypan B 
rob 
ue Uptake 
- 
Noransto IL .PL jjG Hjfjexia 
Hypoxia 
Viability 
f H] Thymidine Uptake 
Hypoxia 
Hypoxia 
Fig. 2. Viability of hypoxic cardiocytes treated with Immunoliposomes (IL), plain liposomes 
(PL), IgG-liposomes (IgG-L) and normoxic and hypoxic conditions determined by 
trypan blue dye exclusion (left panel) or with Tritiated thymidine uptake criteria (right 
panel). 
Delivery of Nanoparticles to the Cardiovascular System 487 
by the dye exclusion method was higher than by [3H]thymidine uptake assessment 
(Fig. 2 right panel). Although the pattern is similar, the absolute difference may be 
due to the less stringent approach for the assessment viability by the trypan blue 
method. 
Inclusion of rhodamine labeled lipids into the formulation of the immunoliposomes, 
enabled visualization of the attachment of liposomes on the surface of 
hypoxic cardiocytes in culture by epifluorescence (Fig. 3) or by confocal microscopy 
(Fig. 4). Only hypoxic cells treated with rhodamine liposomes showed epifluorescence 
(Fig. 3, left), whereas PL treated cells showed no epifluorescence (Fig. 3, 
right). Similarly, confocal micrographs showed that there were discrete regions 
of epifluorescence, the diameter of which corresponded to those of the liposomes 
(~ 200-280 nm).22 Assessment of the number of intact immunoliposomes 
Fig. 3. Epifluorescent micrographs of hypoxic H9C2 cardiocytes treated with rhodamineantimyosin 
immunoliposomes (left) and rhodamine-plain liposomes (right). 
Fig. 4. Black and white confocal micrograph showing localization of intact liposomes 
(left). Pseudocolor of another confocal micrograph showing a pink hue underlying structures 
which appear to be individual liposomes. The bars represent 10 /xm. 
488 Khaw 
on hypoxic cardiocytes with normal cell morphology resulted in 80 ± 20 liposomes 
per cardiocytes (Fig. 4, left). However, there also appears to be diffused fluorescence 
in the cell membrane, indicative of the incorporation of the fluorescent lipids 
from the liposomes into the cell membrane (Fig. 4, right), probably resulting from 
the fusion of the immunoliposome membrane with that of the cell membrane. The 
incorporation of the fluorescent lipid from the immunoliposomes to the cardiocytes 
is not due to the action of lipid transferases, since there are no transferases in the 
culture medium. However, in in vivo situations, such transfer of liposomal lipids to 
normal cell membrane lipid bilayer could occur. 
Preservation of myocardial viability by cell membrane lesion sealing with CSILs 
was also feasible in adult intact hearts.23 In this study, immunoliposomes and control 
liposomes had an average diameter of 200 ± 35 nm. Isolated adult rat (CD-I) 
hearts were perfused with oxygenated Krebs-Henseleit bicarbonate buffer at 37°C, 
after instrumentation on a Langendorff perfusion apparatus.23 Hearts were perfused 
under constant pressure of 80 mm Hg. Each heart was immersed in 0.9% 
NaCl solution maintained at 37° C and was paced at 300 beats/min (5 Hz). The left 
ventricular end-diastolic pressure was set at 10 mm Hg, utilizing a water-filled 
balloon-tipped catheter attached to a pressure transducer. The baseline hemodynamic 
parameters were recorded using a strip-chart recorder after 10 min of 
stabilization period. Global ischemia was induced by decreasing the perfusion pressure 
to zero within 60 seconds. Then, a 2 ml aliquot of freshly prepared 1 mg NGPE 
modified antimyosin immunoliposomes (CSILs), 1 mg NGPE modified non-specific 
IgG-liposomes (IgG-L) or placebo (PBS) was infused at various times of global 
ischemia. Various preparations of liposomes or placebo were administered into 
the aorta via a three-way stopcock placed 8 cm above the aorta, enabling administration 
of various agents without turning on the perfusion pump. This process 
enabled maintenance of global ischemia for the duration of the ischemic period. In 
all studies, a total global ischemia was maintained for 25 min followed by reperfusion 
for an additional 30 min. During the reperfusion period, the end systolic and 
end-diastolic pressures were determined and the difference represented as left ventricular 
developed pressure (LVDP). LVDP of each heart was then compared with 
the baseline LVDP and % LVDP of the baseline was determined. When globally 
ischemic hearts were treated with CSILs at 1 min of ischemia, the recovery of function 
(mean LVDP = 98 ± 14%) during reperfusion was near normal LVDP of sham 
operated hearts (p = NS) (Fig. 5, left), and was highly and significantly greater than 
the LVDP of hearts treated with placebo (12 ± 7%,p = 0.01). The total time function 
curve of the LVDP of hearts treated with CSIL at 1 min of global ischemia was 
87 ± 6% (p = ns versus sham LVDP), but was greater than that of placebo treated 
hearts (12 ± 2%, p = 0.01). Injury to hearts after 25 min of global ischemia that were 
treated with CSIL or placebo, compared with sham operated heart by histochemical 
Delivery of Nanoparticles to the Cardiovascular System 489 
Fig. 5. LVDP of globally ischemic or normal hearts treated with CSIL (•), sham operation (•) 
and placebo (o) during reperfusion for 30 min (left panel), and the corresponding nitroblue 
tetrazolium chloride stained heart slices showing normal myocardium (stained brown) and 
infarcted myocardium (no staining, light color). 
staining with nitroblue tetrazolium, is shown in Fig. 5, right. Nitroblue tetrazolium 
stains for dehydrogenase enzyme activity and is seen as brown to purplish brown 
stained tissues. These enzymes are lost following myocardial or cellular necrosis, 
resulting in no staining of the infarcted tissues seen in Fig. 5 (right), as light colored 
regions in the placebo treated hearts. Quantitative assessment also demonstrated 
that the size of the injury of CSIL treated hearts (4 ± 1% of total ventricles) was the 
same as that of the sham operated hearts (3 ± 2%,p = ns). 
If interventions were instituted almost immediately after the onset of global 
ischemia, then preservation of structure and function of the myocardium would 
be 100%. However, in real-life situations, time of initiation of injury to intervention 
cannot possibly be that short in most circumstances. Therefore, studies were also 
undertaken to determine whether there is a time dependency on myocardial function 
and structural preservation relative to CSIL administration. Thus, Langendorff 
instrumented hearts were subjected to global ischemia as before, however, administration 
of CSIL or control non-specific IgG-L was instituted at 5,10 and 20 min of 
global ischemia. Reperfusion was instituted at 25 min and reperfusion sustained for 
30 min. In globally ischemic hearts with CSIL administration at 5 min of ischemia, 
return to near normal LVDP was achieved at 10 min of reperfusion (Fig. 6, top left 
panel); when CSIL was administered at 10 min of global ischemia, return to near 
normal function was at 15 min of reperfusion (Fig. 6, top right panel). However, 
when CSIL was administered at 20 min of global ischemia, recovery of function 
was only 50 ± 7% of baseline LVDP (Fig. 6, bottom left panel), which was still 
greater than the LVDP of hearts injected with IgG-L (29 ± 5%,p = 0.01). Yet, the 
mean LVDP of all hearts treated with CSIL was statistically greater than the LVDP 
of hearts treated with non-specific IgG-L at corresponding times [Fig. 6 (bottom 
right panel)]. Infarct sizes determined by computer planimetry of the nitroblue 
490 Khaw 
Orran 5rrin 10min 15min 20min 25min 30rrin 
Time of Reperfusion (min) 
-Sham 
-CSILS 
-IgG-L 5' 
(%) LVDP Mean 
90 
so 
70 
60- 
50- 
40- 
30- 
20- 
10- 
n 
Smln 10m In 15mtn 20m In 25mln 30mln 
Time of Reperfusion (min) 
-CSIL 10' 
-IgG-L 10' 
100 i 
90- 
80' 
70 
60 
50' 
40 
30 
20 
10- 
A ^ ^ - 
Omin 5min 10nin 15min 20mn 25nin 30nin 
Time of Reperfusion (min) 
-Sham 
-CSIL 20 
-fcG-L23 
-FBS CSIL 5' CSIL 111' CSIL201 IgG-L 5' IgG-L W IgG-L 2ff 
Treatments 
Fig. 6. LVDP of globally ischemic adult rat hearts treated with CSIL (•), IgG-L (A) or 
placebo (o) relative to sham instrumented control hearts (•) at 5 (top left panel), 10 (top right) 
and 20 min (bottom left panel). Mean LVDP of hearts from 20 to 30 min of reperfusion are 
shown in the bottom right panel. 
tetrazolium chloride stained heart slices showed that hearts treated with CSIL at 1, 
5 and 10 min had similar injury as that of the sham operated hearts (4 ± 1%, 8 ± 3% 
and 6 ± 2%, and 3 ± 2% respectively. P — NS). The infarct size of hearts treated 
with CSIL even at 20 min of global ischemia was 19 ± 3% of the ventricles. This 
was significantly smaller than its corresponding control (p < 0.05), whereas hearts 
treated with control IgG-L at 5, 10 and 20 min of global ischemia were 39 ± 4%, 
35 ± 7% and 45 ± 6% respectively (Fig. 7, left panel). The corresponding nitroblue 
tetrazolium chloride stained heart slices are shown in Fig. 7, right panel. 
Another parameter of myocardial injury that was determined was mitochondrial 
size. Although mitochondrial swelling is a hallmark of ischemic injury, 
irreversible injury cannot be directly extrapolated from just observation of 
mitochondrial size. Nevertheless, in view of the myocardial functional and histochemical 
evidences, mitochondrial size assessment from transmission electron 
micrographs add additional support for myocardial preservation in CSIL treated 
hearts, relative to IgG-L or placebo treated hearts. Figure 8 shows the comparison of 
Delivery of Nanoparticles to the Cardiovascular System 491 
SHAM CSiL CSIL CSSL CSIL IgG-L igG-L IgG-L PBS 
1MN SRfliN 10M1N 2SMIN 5MR4 1QMIN 20H!N 
Fig. 7. Mean infarct sizes of rat hearts treated with CSIL or IgG-L or placebo at 1, 5, 10 
and 20 min of global ischemia (left panel). The corresponding nitroblue tetrazolium chloride 
stained mid slices of rat hearts treated with CSIL or IgG-L at 5, 10 and 20 min of global 
ischemia. Minimal injury was seen in 5 and 10 min CSIL treated hearts, but injury was evident 
in the 20 min CSIL treated heart slice. Injury is evident in all heart slices treated with IgG-L 
(right panel, bottom two rows). 
3090 
» asoo 
2000 
1 1S80 
& 1B00 
800 
Normal CSIL 1* CSIL 5' CSIL 10' CSIL 20' IcjG-L 5' IgG-L IgG-L Placebo 
10* 20' 
Fig. 8. Mean mitochondrial size of normal, CSIL, IgG-L or placebo treated hearts. Treatment 
was as indicated in the text. 
mitochondrial size of normal hearts, CSIL treatment at 1,5,10 and 20 min of global 
ischemia, as well as IgG-L treated hearts at 5,10 and 20 min of global ischemia and 
with placebo. No difference in mitochondrial size was observed between normal 
myocardium (1441 ± 146 mean number of pixels ± SEM) and myocardium treated 
at 1, 5, 10 and 20 min of global ischemia (1496 ± 103, 1496 ± 66, 1845 ± 147 and 
1504 ± 101 respectively) (p = NS). However, mitochondria of hearts treated with 
492 Khaw 
IgG-L at 5, 10 and 20min of global ischemia or placebo (2294 ± 95, 2387 ± 119, 
2667 ± 37 and 2234 ± 270 respectively) were larger than mitochondria of CSIL 
treated hearts (p < 0.05). These studies showed that myocardial viability preservation 
is not restricted to embryonic cardiocytes in cultures. Adult hearts are also 
amenable to structural and functional preservation, following cell membrane lesion 
sealing in a time-dependent manner during ischemia. This method of cell membrane 
lesion sealing has also been reported to preserve the integrity of vascular 
endothelium with antiactin-immunoliposomes.24 
A question that remains concerning the utility of CSIL is whether immunoliposomes 
can retain the protective functions in the presence of plasma proteins in vivo, 
since experiments have demonstrated that cells in culture and adult hearts perfused 
with non-protein oxygenated buffer were prevented from undergoing myocardial 
necrosis, following cell membrane lesion sealing intervention with cytoskeletalantigen 
specific immunoliposomes. To demonstrate that cell membrane lesion sealing 
also occur in vivo, rabbits with experimental myocardial infarction were used. 
In this study, rabbits were injected with anti-myosin CSIL, plain liposomes or saline 
at the time of left circumflex coronary artery occlusion by intracoronary infusion.25 
The occlusion was kept for 45 min followed by 6 hrs of reperfusion. The hearts were 
excised, sliced into ~5 slices parallel to the short axis and stained with nitroblue 
tetrazolium chloride. The infarct was approximately 5 to 10% of the infarcts of the 
control plain liposome or saline treated rabbit hearts.25 Subsequently, comparison 
to IgG-liposome treated hearts with acute myocardial infarction demonstrated that 
the CSIL treatment resulted in significantly smaller infarct size, as was observed in 
comparison to plain liposome or saline treated hearts. 
Thus, cytoskeletal-antigen specific immunoliposomes, consisting of antimyosin 
or antiactin-immunoliposomes, were demonstrated to be able to preserve cell viability 
and integrity. Its potential utility in the cardiovascular system would be 
enhanced once its efficacy following intravenous delivery has been demonstrated. 
However, the study of Asahi etal. showed that intravenous delivery of the antiactinimmunoliposomes 
enabled preservation of the integrity of the endothelial cells of 
the cerebral vessels.24 
5. CSIL as Targeted Gene or Drug Delivery 
Due to the proposed mechanism of cell membrane lesion sealing, we also proposed 
that if drugs or gene constructs were to be included in the immunoliposomes such 
as CSILs, then these drugs or gene constructs should be delivered directly into the 
cytoplasm (Fig. 9). This route should bypass the endocytic route of drugs or gene 
construct delivery, thereby reducing destruction of the delivered cargo by the lysosomal 
enzymes, after formation of endolysosomes. Using silver grains as model 
Delivery of Nanoparticles to the Cardiovascular System 493 
CStwf»DN«, 
:3ajs»flft tigeq^t mmikt *m fesfe* 
Fig. 9. Diagrammatic representation of delivery of intraUposomally entrapped genetic 
construct or drugs directly into the cytoplasm of target cell. 
Fig. 10. Transmission electron micrographs of embryonic cardiocyte treated with 
silver grains impregnated CSIL (left) and plain liposome impregnated with silver grains 
(right). — = 1 /xm. 
drugs, we demonstrated that these drugs can be delivered directly into the cytoplasm 
of hypoxic cardiocytes treated with silver grains loaded CSILs.16'25 Figure 10 
(left) shows a transmission electron micrograph of a cardiocyte treated with silver 
grains impregnated CSILs. Silver grains in groups of concentration at about 
200 nm were observed. However, in cells treated with silver grains impregnated 
plain liposomes, very few cells were viable. Of one such cell detected by transmission 
electron microscopy, the silver grains were observed in the extracellular space 
[Fig. 10 (right)].25 
When the silver grains were replaced with genetic constructs, pGL2 and pSV-^- 
gal vectors, hypoxic cardiocytes treated with CSIL impregnated with either vectors 
showed lucif erase activity or bacterial jS-galactosidase activity. The successful transfection 
of the hypoxic cardiocytes with pGL2, a vector for fire-fly lucif erase enzyme, 
494 Khaw 
3 
& 
** >° ^9 J> c$y « v ^ u ^ 
Fig. 11. Relative light units of luciferase activity of cardiocytes treated with various 
preparations and controls. 
in CSILs is shown in Fig. 11 as relative light units (RLUs). RLUs were determined by 
the use of a luminometer.25 As can be seen, only hypoxic cardiocytes treated with 
pGL2-CSILs showed increased RLUs significantly above normal cells with treatment 
with no vectors. Similarly, normoxic cardiocytes treated with pGL2-CSIL, 
hypoxic cardiocytes and normoxic cardiocytes treated with plain liposomes, or 
with only vectors, showed no significant gene transcription and expression. When 
hypoxic cardiocytes were treated with CSIL with entrapped pSV-/?-gal vectors, 
almost all cells in the field of view under light microscopy exhibited bacterial- 
/6-galactosidase enzyme activity, following reaction with X-Gal (0.2% 5-bromo-4- 
chloro-3-indolyl-beta-D-galactopyranoside, 2nM MgCl2,5 mM K4Fe(CN)6 • 3H20, 
5mM K3Fe(CN)6 in phosphate buffered saline pH 7.4) [Fig. 12(A)]. When this 
mode of gene expression was compared with transfection of pSV-^-gal vector with 
cationic liposomes, cationic liposome transfection according to the manufacturer's 
protocol resulted in transfection of only a few cells per field of view [Fig. 12(B)]. 
In this micrograph, two cells with intense /5-galactosidase activity were observed. 
Quantitation of the number of cells in the field of view that was successfully transfected 
with pSV-/J-gal vector in CSIL, cationic liposome, IgG-liposomes, plain liposome 
and vector alone are shown in Fig. 12(C). Only CSIL and cationic transfection 
showed gene expression. CSIL-transfection or Csilfection was more than 40 times 
more efficient in transfecting cells than cationic liposomes. Although the intensity 
of gene expression was low with Csilfection, using the initial vector concentration 
of 75/Lig of vectors in 13.5 mg lipids in 3 ml, when the vector concentration was 
increased to ~ 200 /xg, also in 13.5 mg lipids in 3 ml, the intensity of gene transfection 
was increased [Fig. 12(D)], This study showed that approximately 3 x 10~12 .ig 
Delivery of Nanoparticles to the Cardiovascular System 495 
*— <——"— 
• 
\ > ' 
« 
- 
« * 
%S . 
A 
* 3 
* - 
- 
B 
CSIL- Cat- CSIL- IgG-L PL- PL- DNAHjp 
Lq> Nor Hyp Hjp Nor Nor 
» < 
* >s 
+ 
Fig. 12. A. Cardiocyte transfection with psV-/i-galactosidase-CSILs at 50 Mg/ml vector concentration. 
Magnification x 400. B. Cationic liposome transfection with psV-/9-galactosidase 
vectors as per package insert. Magnification x 400. C. Number of cells transfected with psV- 
/3-gal vectors by CSIL, cationic liposomes or plain liposomes. D. Csilfection with 140 /xg/ml 
psV-/i-galactosidase vectors. Magnification x 100. 
of DNA were delivered into the cytoplasm of each cardiocytes, whereas cationic 
liposomes delivered approximately 9.5 x 10~6 /xg DNA per cell.25 Yet, by increasing 
the DNA content by 3 times in the CSILs, the intensity of /3-galactosidase expression 
was increased to the level of cationic liposome transfection, with at least 40 times 
more cells transfected. 
6. Conclusion 
The application of nanoparticles in the cardiovascular system is finally becoming 
desirable. As this chapter has shown, the initial foray into this field occurred in 
1979, even though the term nanotechnology was not coined for at least a decade. 
Nevertheless, its potential as drug delivery and targeting for therapy and diagnosis 
were recognized earlier on. To date, the application of targeted nanoparticles in the 
cardiovascular system has included targeting the endothelium of atherosclerotic 
lesions and other inflammatory processes, gene delivery to ischemic cardiocytes 
and cell membrane lesion sealing with cytoskeletal-antigen specific immunoliposomes. 
Other applications such as targeted drug release from nanoparticles, after 
496 Khaw 
targeted drug localization in the cardiovascular system, is envisioned for future 
therapy. 
References 
1. Allen TM, Cullis PR (2004) Drug delivery systems: Entering the mainstream. Science 
303:1818-1822. 
2. Hoet PHM, Bruske-Hohlfeld I, Salta OV (2004) Nanoparticles-known and unknown 
health risks. / Nanobiotechnol 2:12. 
3. Hood JD, Bednarski M, Frausto R, Guccione S, Reisfeld R, Xiang and Cheresh DA (2002) 
Tumor regression by targeted gene delivery to the neovasculature. Science 296:2404-2407. 
4. Kralj M and Pavelic K (2003) Medicine on a small scale. Europena Molecular Biology 
Organization Reports 4:1008-1012. 
5. Caride VJ and Zaret BL (1977) Liposome accumulation in regions of experimental 
myocardial infarction. Science 198:735-738. 
6. Torchilin VP, Khaw BA, Smirnov VN and Haber E (1979) Preservation of antimyosin 
antibody activity after covalent coupling to liposomes. Biochem Biophys Res Coram 98: 
1114-1119. 
7. Khaw BA, Torchilin, VP, Berdichevskii VR, Barsukov AA, Klibanov AL, Smirnov VN 
and Haber E (1983) Enhancing specificity and stability of targeted liposomes by coincorporation 
of sialoglycoprotein and antibody on liposomes. Bull Expt Biol Med (Translated 
from Russian). 95:776-781. 
8. Klibanov AL, Maruyama K, Torchilin VP and Huan L (1990) Amphipathic polyethyleneglycols 
effectively prolonged the circulation time of liposomes. FEBS Lett 268: 
235-237. 
9. Torchilin VP, Klibanov AL, Huang L, O'Donnell S, Nossiff ND and Khaw BA (1992) 
Targeted accumulation of PEG-coated immunoliposomes in infarcted myocardium in 
rabbits. FASEB 6:2716-2719. 
10. Torchilin VP, Narula J, Halpern E and Khaw BA (1996) Poly (ethylene glycoD-coated 
anti-cardiac myosin immunoliposomes: Factors influencing targeted accumulation in 
the infarcted myocardium. Biochim Biophys Acta 1279:75-83. 
11. Lanza GM, YU X, Winter PM, Abendschein DR, Karukstis KK, Scott MJ, Chinen LK, 
Fuhrhop RW, Scherrer DE and Wickline SA (2002) Targeted antiproliferative drug delivery 
to vascular smooth muscle cellls with a magnetic resonance imaging nanoparticle 
contrast agent. Circulation 106:2842-2847. 
12. Lanza GM, Wallace KD, Scott MJ, Cacheris WP, Abendschein DR, Christy DH, 
Sharkey AM, Miller JG, Gaffney PJ and Wickline SA (1996) A novel site-targeted ultrasonic 
contrast agent with broad biomedical application. Circulation 94:3334-3340. 
13. Spragg DD, Alford DR, Greferath R, Larsen CE, Lee KD, Gurther GC, Cybulsky MI, 
Tosi PF, Nicolau C and Gimbrone Jr MA (1997) Immunotargeting of liposomes to activated 
vascular endothelial cells: A strategy for site-selective delivery in the cardiovascular 
system. Proc Natl Acad Sci USA 94:8795-8800. 
Delivery of Nanoparticles to the Cardiovascular System 497 
14. Bloeman PG, Henricks PA, van Bloois L, van den Tweel MC, Bloem AC, 
Nijkamp FP, Crommelin DJ and Strom G (1995) Adhesion molecules: A new target for 
immunoliposome-mediated drug delivery. FEBS Lett 357:140-144. 
15. Hamilton AJ, Huang SL, Warnick D, Rabbat M, Kane B, Nagaraj A, Klegerman M and 
McPherson DD (2004) Intravascular ultrasound molecular imaging of atheroma components 
in vivo. J Am Coll Cardiol 43:453^160. 
16. Khaw BA, Torchilin VP, Vural I and Narula J (1995) Plug and seal: Prevention of hypoxic 
cell death by sealing membrane lesions with cytoskeleton-specific immunoliposomes. 
Nat Med 1:1195-1198. 
17. Shi R, Qiao X, Emerson N and Malcom A (2001) Dimethylfulfoxide enhances CNS neuronal 
plasma membrane resealing after injury in low temperature or low calcium. JNeurocytol 
30(9-10):829-839. 
18. McNeil PL (2002) Repairing a torn cell surface: Make way, lysosomes to the rescue. / Cell 
Sci 115(Pt 5):873-879. 
19. Togo T, Alderton JM and Steinhardt RA (2003) Long-term potentiation of exocytosis and 
cell membrane repair in fibroblasts. Mol Biol Cell 14:93-106. 
20. McNeil PL and Ito S (1989) Gastrointestinal cell plasma membrane wounding and resealing 
in vivo. Gastroenterology 96:1238-1248. 
21. Walev I, Hombach M, Bobkiewicz W, Fenske D, Bhakdi S and Husmann M (2002) Resealingoflarge 
transmembrane pores produced by streptolysin O in nucleated cells is accompanied 
by NF-kappa B activation and downstream events. FASEB 16(2):237-239. 
22. Khaw B A, da Silva J, Vural I, Narula J, Torchilin VP (2001) Intracy toplasmic gene delivery 
for in vitro transfection with cytoskeleton-specific immunoliposomes. / Control Rel 75: 
199-210. 
23. Khudairi T and Khaw BA (2004) Preservation of ischemic myocardial function and 
integrity with targeted cytoskeleton-specific immunoliposomes. / Amer College Cardiol 
4:1683-1689. 
24. Asahi M, Rammohan R, Sumii T, Wang X, Pauw RJ, Weissig V, Torchilin VP and Lo EH 
(2003) Antiactin-targeted immunoliposomes ameliorate tissue plasminogen activatorinduced 
hemorrhage after focal embolic stroke. / Cerebral Blood Flow Metabolism 8: 
895-899. 
25. Khaw BA, Vural I, Da Silva J, Torchilin VP (2000) Use of cytoskeleton-specific immunoliposomes 
for preservation of cell viability and gene delivery. STP Pharma Sciences 
10(4):279-283. 
This page is intentionally left blank
23 
Nanocarriers for the Vascular 
Delivery of Drugs to the Lungs 
Thomas Dziubla and Vladimir Muzykantov 
The lungs perform a vital multifunctional physiological role. Yet, the pulmonary 
vasculature is susceptible to a host of pathologies, which contribute to morbidity 
and mortality. Many medical interventions can improve the course and outcome 
of these disease conditions, provided they can be delivered in an effective, 
localized and safe manner. Venous administration is a suitable route for drug delivery 
to the pulmonary vasculature, but most drugs do not have the pharmacokinetic 
features required for optimal pulmonary delivery. In theory, this problem 
may be overcome through the use of nanocarriers, which can act to improve the 
localization of drugs in the pulmonary vasculature and allow for a more controled 
release/activity profile for drugs that are otherwise cleared or inactivated rapidly. 
Several types of nanocarriers are potentially useful for this purpose including protein 
conjugates, liposomes and polymer nanocarriers. Stealth coats improve carrier 
circulation, while affinity ligands provide targeting. Yet, despite these promises 
and many experimental advances, significant obstacles must be overcome to permit 
clinical utility. This chapter gives a background of the biomedical aspects of 
lung targeting, introduces basic elements of current design of systems for vascular 
drug delivery to the lungs, and discusses specific applications where nanocarriers 
can improve current therapies, as well as the limitations of existing nanocarrier 
technologies in this context. 
499 
500 Dziubla & Muzykantov 
1. Introduction 
Due to its critical, diverse physiological roles and high vulnerability to pathological 
processes, the pulmonary vasculature represents an important pharmacological 
target. In order to manage lung pathologies, a plethora of diagnostic and therapeutic 
treatments including contrast agents, isotopes, anti-inflammatory, anti-thrombotic 
and antioxidant agents, anticancer and anti-proliferative agents, enzyme replacement 
therapies (ERT), has been proposed. Yet, due to unfavorable natural pharmacokinetic 
properties, many of these strategies are currently not in use. For instance, 
despite the diversity of the chemical classes of these therapeutic agents, many of 
which are bio-therapeutics, e.g. proteins, most of them do not naturally accumulate 
in the lungs after intravascular injection, thereby greatly limiting their effectiveness 
and specificity1 
Many of these limitations may be overcome by the use of nanocarriers, which 
can improve drug delivery to the therapeutic site by passive and active targeting. 
Furthermore, nanocarriers can optimize the pharmacokinetic properties of drugs 
by: (1) increasing the delivery potential of poorly water-soluble drugs (2) providing 
extended release of drug in localized areas (3) enhancing the circulation life-time 
and (4) isolating sensitive /bioactive drugs from the blood, protecting from premature 
inactivation and systemic adverse effects. This chapter focuses on nanocarriers 
designed for drug delivery to the pulmonary vasculature. It begins with a brief 
background of the lungs as a therapeutic target and describes nanocarriers design, 
potential applications, current limitations and avenues for optimization and translation 
into the clinical domain. 
2. Biomedical Aspects of Drug Delivery to Pulmonary 
Vasculature 
While gas exchange, providing blood oxygenation in the vascular system, is the 
most important pulmonary function, the lungs serve a variety of other vital 
functions.2 For instance, the pulmonary vasculature, a unique anatomical and functional 
compartment itself, acts as an anatomical filter for thrombi, aggregates activated 
or damaged blood cells and other types of emboli (e.g. lipid, gas) in the venous 
blood, which would otherwise embolize cerebral vasculature, resulting in stroke. 
In addition, with enzymes exposed on the luminal side of the vascular walls, it 
functions as a reactor bed for the blood, converting circulating agents (e.g. peptides, 
mediators and hormones), and thereby affecting systemic signaling and physiology. 
The pulmonary vasculature is the primary interface between the systemic circulation 
and the exterior environment. Hence, it is vulnerable to the damaging effects 
of extraneous (e.g. inhaled pollutants, particulates, and pathogens) and endogenous 
Delivery of Nanocarriers to the Lungs 501 
pathological factors (e.g. circulating thrombi, pathogens, tumor metastases). In particular, 
the pulmonary vascular endothelium, a cellular monolayer lining the luminal 
surface of blood vessels, is involved in many pathological conditions, local and 
systemic, and represents an important target for diverse diagnostic, prophylactic 
and therapeutic interventions. This chapter will focus on advanced drug delivery 
systems designed to achieve this specific goal. 
2.1. Routes for pulmonary drug delivery: Intratracheal vs vascular 
Drugs can easily reach lung tissue through either intratracheal (IT) or intravenous 
(IV) administration. Upon considering pulmonary drug delivery, IT administration 
(e.g. aerosols, inhalants) is the first to come to mind. It provides a route for noninvasive 
means of drug delivery to airway compartments (e.g. bronchial epithelium 
and interstitium)3 and beyond, into the systemic circulation. As such, this is an 
ideal situation for drugs such as asthma medications, where bronchiolar delivery 
is required, or for systemic delivery of drugs (e.g. hormones), which can pass via 
the epithelial cells and other components of gas-blood barrier.4-6 
However, for diseases where delivery to pulmonary endothelial cells is needed, 
IT administration effectiveness is limited. This route provides patchy delivery with 
inconsistent alveolar reach.7 Since the alveoli are the area of greatest vasculature 
density with slowest perfusion, it is the key, yet relatively difficult to reach, site for 
the transport of drugs from the airways to circulation. Furthermore, once transport 
to the vascular space does occur, nothing keeps drugs from fleeing into the systemic 
circulation, thereby resulting in insufficient local residence time and concentration, 
thus limiting therapeutic effects in the pulmonary endothelium. 
In contrast to the intratracheal route, IV is naturally designed to aid the delivery 
of circulating compounds to the pulmonary endothelium. The pulmonary vasculature 
is the first major microvascular network, which represents one third of the 
entire vascular surface area, encountered by IV injected drugs. In addition, the lungs 
receive half of the cardiac output at each systole (i.e. entire venous blood), whereas 
all the other organs share the other half (i.e. arterial blood). Also, the rate of blood 
perfusion through the high-capacity, low pressure vascular system in the lungs is 
relatively slow (see below), favoring interactions of circulating ligands with pulmonary 
endothelium. For these reasons, this review will focus on vascular targeting 
of the pulmonary endothelium by IV injection. 
2.2. Pulmonary vasculature as a target for drug delivery 
To design systems for pulmonary vascular drug delivery, one has to know pertinent 
features of lung vasculature physiology. For example, in order to cope 
502 Dziubla & Muzykantov 
with bouts of high cardiac blood output and to satisfy oxygen demands, the 
lungs possess a high transient perfusion capacity. Hence, blood pressure and the 
rate of perfusion in the lungs are significantly lower, compared with systemic 
vasculature. 
Several mechanisms regulate perfusion via pulmonary vasculature to adjust 
to changing cardiac blood output and ventilation rate. The lower lobes of both left 
and right lungs are perfused more effectively than the apical lobes. This inequity 
is matched by a similar ventilation pattern that exists between basal and apical 
areas of the lungs. Hence, lower lobes will receive more injected or inhaled drugs 
(Fig. 1). 
A substantial fraction of pulmonary capillaries are only transiently perfused 
and they get recruited in physical stress for a greater blood volume exchange, 
optimizing the rate of gas exchange. These transiently perfused vessels, forming 
a reserve perfusion capacity to suit physical stress, can also be recruited to cope 
with the redistribution of blood flow in cases of localized and systemic pathologies 
(e.g. heart diseases). For instance, when vessels are partially or fully occluded 
(e.g. by fibrin thrombi or activated white blood cells), the adjacent vasculature is 
Distribution of Drug Delivery 
Areas of Inflammation/ 
Enhanced Drug Localization 
Fig. 1. Normal and pathological pulmonary blood perfusion patterns affect distribution 
of delivered nanoparticles. Under normal conditions, due to preferential perfusion and ventilation 
of lower lobes, this area of lungs will accumulate higher loads of nanocarriers (left). 
Enhanced permeability of pulmonary vasculature will drive preferential delivery of nanoparticles 
to sites of acute inflammation, hence passive targeting (right). Areas of inflammation 
may be reached by EPR effects or by active targeting of nanocarriers with coated with antibodies 
to cell adhesion molecules expressed preferentially in areas of inflammation. This 
allows for both the treatment of lung inflammation and non-invasive visualization areas of 
lung pathology. 
Delivery of Nanocarriers to the Lungs 503 
recruited to meet perfusion demand and compensate for the pathological deficit. 
This ability to rapidly respond and alter flow patterns, allows the lungs to function 
as a filter for debris that would otherwise embolize the brain and the other 
organs. 
Pulmonary perfusion changes under pathological conditions, thus affecting 
drug delivery. In addition to focal perfusion changes caused by thrombosis or 
inflammation (Fig. 1), generalized pulmonary vascular pathologies markedly alter 
hemodynamics in this organ. For example, primary pulmonary hypertension, 
depending on the phase of the disease, might lead to either acceleration or deceleration 
of pulmonary perfusion. Congestive heart disease, defects of the right heart 
valves and the insufficiency of pumping function of the right ventricle, may all 
result in blood pooling and stagnation in the pulmonary vasculature. All these 
factors may affect pulmonary delivery of injected drugs. 
3. Pulmonary Targeting of Nanocarriers 
Selective localization of drugs in the site of interest can be achieved by passive 
and/or active targeting. Passive targeting refers to the accumulation of carriers not 
involving specific recognition of the target compartment in the body, and includes 
mechanical and charge retention, and the enhanced permeation and retention (EPR) 
effect. In most cases, active targeting that employs recognition moieties possessing 
specific affinity to target determinants (e.g. antigen-antibody8-10 or receptorligand,
11-14) affords greater specificity of drug delivery. This section reviews these 
strategies for delivery nanocarriers to the pulmonary vasculature. 
3.1. Effects of carrier size on circulation and tissue distribution 
Whether passive or active targeting is used, nanocarrier size can affect its distribution, 
circulation and subcellular localization (Fig. 2). When carrier size is < 100 nm, 
permeation across endothelial and epithelial barriers is possible via transcellular 
and peri-cellular pathways.15 Sub-micron carriers are less likely to pass through 
intercellular junctions in endothelial and epithelial cells, with the exception of 
organs with fenestrated endothelium having large, few micron openings, such as 
in the liver and the spleen. However, even relatively large carriers of ~500nm 
in diameter, are still capable of being internalized either via receptor-mediated 
(e.g. endocytosis) or constitutive (e.g. macropinocytosis) pathways. Cellular internalization 
allows for a more precise level of control of subcellular destinations 
including lysosomes, other intracellular compartments and cytosol, or even beyond 
the endothelial cells.15 
Micron carriers still allow circulation without embolism, although the likelihood 
of either barrier penetration or cellular internalization is greatly limited. 
504 Dziubla & Muzykantov 
. » ' . • m m 
<100nm < 500 nm >1000nm 
Fig. 2. Effect of carrier size on transport through vascular endothelium. Nanocarriers 
< 100 nm diameter are capable of passing though certain endothelial barrier either between 
the cells or via transcellular mechanisms involving endocytosis. Nanocarriers < 500 nm 
poorly transport between endothelial cells in the lungs, yet they are still capable of being 
internalized by endothelium. Particles larger than 1 /xm, may still be capable of circulating 
and being targeted, yet they are unlikely to leave vascular lumen in the lung unless pathological 
factors induce abnormally high vascular permeability (leakiness, not shown). 
Such size ranges provide a mechanism for maintaining targeted drug carriers to 
reside on the luminal side of the endothelium, an ideal situation for drugs that 
require blood/plasma contact for therapeutic activity.16 
Size also determines the carrier's fate in the circulation. Despite the fact that 
sub-micron size range permits unimpeded vascular circulation, nanocarriers are 
cleared from the bloodstream within minutes via uptake by reticuloendothelial 
system (i.e. RES, including hepatic and spleenic resident macrophages available to 
the blood, via openings in the vessels). In mice, this can result in 60-90% clearance 
of the injected dose in the first instance.17-19 
Grafting the surface of nanocarriers with large molecular weight hydrophilic 
polymers, negative or neutral, the primary example being poly(ethylene glycol) 
(PEG), greatly extends the circulation time.17'20 PEG modified carriers (stealth) 
have a hydrophilic molecular brush that repels cellular and protein interactions, 
thus reducing recognition and uptake by RES.21 Tissue uptake of PEG-coated 
carriers depends more on mechanical retention than on active recognition and 
phagocytosis by RES; hence smaller, carriers circulate for longer duration than 
large ones. 
There is growing evidence that carrier geometry is critical to circulation and 
cellular localization effects. For instance, worm-like micelles have been reported to 
align with flow, a feature that has been hypothesized to extend and prolong the 
circulation of stealth carriers.22 Also, liposomes containing polymerized micelles 
possessed both an elongated, ellipsoidal shape, as well as a greatly enhanced circulation 
profile.23,24 It is not clear whether these effects are a result of improved fluid 
dynamics or phagocytic evasiveness. However, this effect allows for additional levels 
of design/control of circulation, and perhaps other pharmacokinetic features of 
nanocarrier systems. 
Delivery of Nanocarriers to the Lungs 505 
3.2. Passive targeting 
3.2.1. Mechanical retention 
Microspheres larger than the pre-capillaries (i.e. >10 micron diameter) injected 
into the venous system, embolize the downstream capillary bed. Thus, the site 
of injection dictates the localization site; hence, targeting lung vasculature can be 
achieved by simply injecting into a pulmonary artery or an upstream femoral vein. 
Since embolism occurs at the first bifurcation that is too small for carrier passage, 
targeting is limited to the arterioles. Delivery to the venous sites occurs 
only in a form of released drug passage through this downstream vascular compartment. 
Furthermore, while pulmonary vasculature can tolerate low levels of 
embolism, it is not a fully benign process, resulting in ischemic vascular pockets 
losing contact with the blood flow and the nutrient exchange. Yet, mechanical 
retention of degradable microcarriers in the pulmonary vasculature has medical 
utility, e.g. for visualization of the lung blood vessels and perfusion patterns using 
radiolabeled microspheres. Furthermore, newer treatments for massive hemoptysis 
(the coughing up of blood) have focused on the embolization of the bronchial 
arteries.25 
While microspheres embolize vasculature, nanocarriers' size allows for 
unobstructed flow throughout all vessels. Yet, nanocarriers can also be designed 
to associate into micron-sized aggregates, prior to or upon injection, which are 
then delivered and mechanically retained in the capillary bed (Fig. 3). Through the 
proper selection of nanocarrier size and rate of aggregate breakup, either subcellular 
or transcellular compartments can be reached, disconnecting embolism and 
drug delivery, and allowing for shorter durations of ischemia with a longer term 
drug delivery phase. 
Further, < 500 nm diameter nanocarriers may provide a more favorable 
degradation pattern, compared with solid microspheres degrading via either surface 
erosion or bulk degradation. For a more detailed discussion, see the reviews 
at Refs. 26-28. Since surface erosion results in the overall shrinking of a particle, 
the remnant microspheres will eventually be washed away from the delivery site, 
prematurely terminating local effects. Bulk degradation is more suitable for a stable 
deposition of microspheres, since the overall structure remains intact until the 
polymer has degraded to the point where structural integrity is completely lost. 
Yet, under a continuous back pressure in the vasculature, particle disintegration can 
result in highly disordered debris of various sizes, geometry and surface roughness 
that can induce local and systemic damage. In the case of aggregated nanocarriers, 
such hazardous debris is likely to be avoided, since individually released nanocarriers 
possess designed nano-scale geometry, permitting non-obtrusive behavior in 
the circulation. 
506 Dziubla & Muzykanlov 
Fig. 3. Mechanical retention of nanocarriers in the pulmonary vasculature. (A) In the 
presence of cross-linking stimuli (e.g. plasma opsonins or circulating ligands in blood), large 
(~ 10-50 nm) aggregates of nanocarriers will form after injection and embolize the pulmonary 
capillary bed, thus creating a high local concentration of a drug and ceasing blood flow. 
(B) As the aggregate disintegrates, released individual nanocarriers can diffuse into the surrounding 
tissue via inter-endothelial gaps or/and transcellular pathways, allowing them to 
accumulate in the pulmonary interstitium. Disintegration of emboli initiates repcrfusion of 
blood. (C) As disintegration proceeds and blood flow is reestablished, released nanocarriers 
will be washed away. Drugs delivered by and released from aggregated nanoparticles will 
be eliminated by the restored flow. 
3.2.2. Charge-mediated retention and non-viral gene delivery 
Nanocarriers possessing a positive surface charge accumulate in the first vascular 
bed, similar to the targeting behavior of mechanical retention, although the mechanism 
of retention is different. The highly anionic glycocalyx covering the endothelium 
binds cationic molecules and particles.29-31 In cell cultures, such binding has 
resulted in the internalization and enhanced levels of transfection by non-viral 
DNA delivery means, e.g. cationic liposomes.32 Yet, many blood components are 
also negatively charged. Hence, the aggregation of serum components and/or the 
thrombus formation resulting in embolism may also occur.33 
High levels of lung targeting due to charge retention in the pulmonary vasculature 
have been displayed by IV injected cationic liposomes and carriers decorated 
with either cationic polymers (e.g. polylysine) or peptides (e.g. TAT) sequences.30,34 
Delivery of Nanocarriers to the Lungs 507 
While it is not clear if in vivo localization is due to particle-endothelium association 
or aggregation, it does provide an interesting mechanism for the internalization 
and cytosolic delivery of DNA for gene delivery. Interestingly, in many instances, 
charge-mediated retention of the non-viral gene delivery means in the pulmonary 
vasculature results in transgene expression in cells underlying endothelium (e.g. 
vascular smooth muscle cells), but not in endothelial cells per se.35 
3.2.3. Pulmonary enhanced permeation-retention (EPR) effect 
The enhanced permeation and retention effect was originally described when long 
circulating stealth liposomes were found to accumulate into vascularized solid 
tumors, due to the erratic, highly permeable nature of the tumor vasculature.36,37 
As nanocarriers circulate and encounter this area, characterized also by poor lymphatic 
drainage, leakage into and retention in the interstitium resulted in gradual 
accumulation. EPR targeting improved with increased circulation times, and when 
nanocarrier size is small enough to pass through the pores in the leaky vessels of 
<200nm. 
A similar mechanism has been found to enhance the delivery into the sites 
of inflammation, where the vasculature is also highly permeable.38 Since the pulmonary 
vascular bed receives the entire venous blood flow and is highly susceptible 
to enhanced vascular permeability under pathological conditions, it is plausible that 
EPR-related accumulation in the lungs might occur. This mechanism might permit 
the visualization of inflammation sites in the lungs and provide a means of treating 
localized pulmonary inflammation and edema (Fig. 1). 
3.3. Active targeting 
Active targeting involves the engagement of specific recognition ligands with surface 
determinants present in the site of interest. This can be achieved by either 
using immunoglobulins raised against target antigens, affinity peptides or using 
a native ligand receptor pair. For a review of endothelial determinants used as 
targets and antibodies, and other affinity ligands used as vectors for active drug 
targeting into the pulmonary vasculature, please see reviews at Refs. 8 and 39. A 
brief list of the key guidelines in pulmonary target selection includes the following 
factors: 
(1) The target should be present on the luminal surface of pulmonary endothelium, 
accessible spatially and temporally, and should not be down regulated 
or masked in disease states. For example, adhesion of activated blood cells 
and accelerated shedding inhibit targeting to some constitutive endothelial 
determinants.40 On the other hand, determinants exposed on the endothelial 
508 Dziubla & Muzykantov 
cells under pathological conditions (e.g. selectins) have a distinct transient surface 
expression profile, which may permit selective drug delivery to pathologically 
altered endothelium, but require exact timing of administration to match 
the duration of target availability. 
(2) The target should not be present in non-endothelial counterparts that are accessible 
to the circulating nanocarriers. For example, endothelial cells have transferrin 
receptors, which are also abundantly exposed in hepatic cells that are 
accessible to the bloodstream. As a result, transferrin-targeted drugs accumulate 
in the liver with minimal delivery to the lungs. Also, analogues of the 
target determinants circulating in the blood (e.g. soluble forms of transmembrane 
glycoproteins or P-selectin on platelets) will compete with endothelial 
counterparts, compromising targeting. 
(3) Targeting should not cause harmful side effects in the vasculature. Binding of 
targeted drugs may cause shedding, internalization, or inhibition of endothelial 
determinants, which may be detrimental. For example, thrombomodulin, a surface 
protein responsible for thrombosis containment, is abundantly expressed in 
the pulmonary vasculature, providing high pulmonary targeting specificity.41 
Yet, its inhibition by antibodies may provoke incidences of thrombosis that 
prevents clinical potential for drug delivery. Ideally, engaging of the target 
should provide therapeutic benefits such as the inhibition of pro-inflammatory 
molecules. 
(4) It is ideal for the docking to a surface determinant to result in optimal subcellular 
addressing of a drug.15 Thus, depending on the therapeutic goal, 
a targeted nanocarrier should either be retained on the cell surface (blood 
therapies) or undergo trafficking to a proper sub-cellular compartment (e.g. 
nucleus in the case of DNA,41 or lysosomes in the case of enzyme replacement 
therapies42). 
No single targeting suits all therapeutic needs. Specific therapeutic goals require 
different secondary effects mediated by binding to the endothelium, drug targeting 
to different sub-populations of endothelial cells, and diversifying the cellular 
compartments. A plethora of affinity carriers, sometimes directed to relatively similar 
endothelial targets (e.g. cell adhesion molecules) or even binding to different 
domains of the same target molecule, are currently explored to capitalize more fully 
on unique opportunities offered by vascular targeting.8,39'43 
Strategies for defining molecular determinants (targets) for affinity delivery 
of nanocarriers to endothelial cells, include both high-throughput analyses 
(e.g. in vivo selection of phage display libraries,43 subtractive proteomics 
of endothelial plasma membrane39) and low-throughput analysis of affinity ligands 
to identify endothelial molecules with known functions.44 Some of the most 
Delivery of Nanocarriers to the Lungs 509 
promising endothelial determinants for such ligands include constitutive antigens 
such as angiotensin-converting enzyme (ACE),44""46 cell adhesion molecules 
of Ig-superfamily (PECAM and ICAM),16'47'48 inducible adhesion molecules 
(E- and P-selectins, VCAM-1),49-53 aminopeptidases and caveoli-associated 
glycoproteins.54-56 
4. Carrier Design 
As a whole, nanocarriers require a "ground up" design approach for each application. 
Depending on the particular needs of a given strategy, material selection can 
vary greatly. This section will outline the general considerations of the design of 
nanocarriers for pulmonary drug delivery. 
4.1. Biocompatibility 
The initial material constraint is biocompatibility, a term that might be misleading, 
without considering the context of a given application. The materials used 
for nanocarriers should induce no deleterious (e.g. thrombogenic, mutagenic or 
carcinogenic) effects in the body. These effects (like with any medicines) depend 
on dose, location, structure, and residence time of nanocarriers. For this reason, 
while pre-labeling a material as "biocompatible" has been used in many papers, it 
provides rather limited information to specific situations and applications. A rigorous 
re-evaluation of carriers' biocompatibility for each new indication in a given 
pathological context (likely, even in given patients cohorts), does not seem to be an 
excessive precaution in a post-Vioxx era. 
For instance, titanium and titanium oxide coated implants has long been 
considered a highly inert, biocompatible material in bone prosthetics and dental 
implants.5758 Yet, sub 100nm nanoparticle forms of titanium oxide have highly 
active surface sites capable of catalyzing the formation of oxygen radicals, which 
can result in cell and tissue injury.59-61 As such, the "biocompatible" label must not 
simply be given to titanium oxide nanoparticles, although this does not mean that 
there is no potential therapeutic use of this carrier. However, there are settings in 
which its use is unadvisable, e.g. drug delivery into the pulmonary tissue which is 
prone to oxidative stress, due to high level of oxygen and reactive oxygen species 
produced by leukocytes and pulmonary endothelial cells.10'62 
On the other hand, some materials that have been previously labeled as nonbiocompatible 
may be revisited for use in nanocarriers, having to undergo degradation 
and excretion pathways unsuitable for larger carriers. However, the primary 
requirement of nanocarrier compatibility is the ability to break down into non-toxic, 
plasma soluble components that can be eliminated via renal filtration or hepatic bile 
510 Dziubla & Muzykantov 
excretion. For this reason, most carriers under development are composed of either 
degradable polymers, or possess MWs lower than 40 KDa.63,64 
4.2. Material selection (by application) 
4.2A. Imaging 
The lungs are a classically difficult organ for imaging due to low-signal to noise 
ratio, multiple air-tissue interfaces, and physiological motion such as cardiac and 
ventilating.65,66 Of all imaging technologies available, the most commonly used 
technology for pulmonary imaging (except routine chest X-rays) is computer 
tomography (CT). Yet, it is still difficult to properly identify many pulmonary disease 
pathologies. The use of targeted contrast agents may allow for the improved 
identification of these disease states.66 In the case of CT, high density materials 
(e.g. metals, crystalline polymers and high atomic weights) are ideal candidates. 
Indeed, early studies using iodinated nanoparticles have been used for the imaging 
of lymph nodes.67,68 
In spite of its utility, CT resolution is limited to ~ 1 mm. NMR, a higher resolution 
imaging technology, has been classically limited to the use of pulmonary 
imaging. Yet, current advancement in imaging algorithms and contrast targeting 
may improve NMR imaging of diseases such as acute pulmonary embolism and 
chronic infiltrative disease.66,69,70 
4.2.2. Gene delivery 
Initial success with gene delivery to the pulmonary tissue was obtained using adenoviral 
carriers. Indeed, heat shock protein HSP70, nitric oxide synthase (NOS), 
and interleukin-10 have all been adenovirally transfected into pulmonary endothelial 
cells, for the attenuation of ischemia-reperfusion injury.71-73 However, systemic 
adenoviral transfection is greatly limited due to an associated cytokine release and 
immune response. In this context, enhancement of local transfection by re-targeting 
viral gene delivery is a highly promising strategy74'75 to pulmonary endothelium 
(e.g. using ACE antibody coupled to viral particles). 
Non-viral gene delivery poses an interesting set of material requirements, 
allowing for the effective delivery of DNA into a target cell and the subsequent 
trafficking of the DNA into the nucleus. These carriers must be able to load high 
levels of DNA into a single particle, and be able to target endothelial cells with the 
subsequent internalization and endosomal escape mechanism to allow for the DNA 
to reach the nucleus. Most of these processes have focused on charge coupling to 
condense DNA into a nanoscale aggregate. The most common of these have been 
the use of cationic polyplexes.76,77 For example, polycationic electrolytes such as 
Delivery of Nanocarriers to the Lungs 511 
poly(ethylenimine) (PEI) and poly(l-lysine) (PLL) have been used to condense the 
negatively charged DNA. PEI (of small chain length) has been shown to reverse 
charge at endosomal pH and release DNA.78-79 
Pulmonary vascular delivery of DNA was possible with cationic surface charge 
alone,80 yet lung specificity can be greatly improved upon application of immunotargeting 
toward endothelial markers such as thrombomodulin,41 PECAM-181 or 
ACE.82-83 
While highly cationic vectors also display a significant inflammatory 
response,84 this immune reaction can be greatly attenuated without a reduction 
in degree of transfection by lowering the overall carrier charge.29'32,85 
4.2.3. Delivery of therapeutic enzymes 
Examples of enzymatic therapies amenable pulmonary targeting using nanocarriers, 
include delivery of: (i) lysosomal enzymes (enzyme replacement therapy, ERT), 
for the treatment of non-neuronal lysosomal storage diseases that affect pulmonary 
endothelium (e.g. Niemann-Pick disease),42 (ii) anti-thrombotic enzymes (e.g. 
plasminogen activators) for the dissolution of blood clots formed or lodged in 
the pulmonary vessels, and (iii) antioxidant enzymes, for the containment of 
vascular oxidative stress in the lungs, which is a highly morbid pathological 
condition. 
Targeting can be achieved by the chemical coupling of enzymes with affinity 
carriers, producing nano-scale protein conjugates. For example, catalase conjugated 
with antibodies to endothelial antigens ACE, PECAM or ICAM, accumulates in the 
lungs of laboratory animals after IV injection and protects against oxidative injury 
in the models of human diseases such as lung transplantation ischemia/reperfusion 
injury86 and acute edematous vascular oxidant stress.87 On the other hand, targeting 
of plasminogen activators to endothelial cell adhesion molecules boosts antithrombotic 
capacity of the pulmonary vasculature.16 Targeting enzymes clearly 
illustrates the importance of proper sub-cellular addressing of drugs, namely, luminal 
surface for fibrinolytics, non-degrading intracellular compartments for antioxidants, 
and lysosomes for ERT.42 
Loading into nanocarriers might optimize some of the enzyme therapies. For 
example, antioxidant catalase loaded into H202-permeable, protease-resistant polymer 
nanocarriers88 might retain its protective activity even within lysosomes. 
Yet, loading into highly amphiphilic carriers (i.e. micelle form, vesicle form) may 
cause undue folding and inactivation of enzyme. Optimally, the carrier material 
would stabilize protein in an anhydrous state to avoid inactivation. This can 
theoretically be achieved via the hydrophobic sequestering of solid protein into a 
polymer core. 
512 Dziubla & Muzykantov 
4.2.4. Sma II molecule drugs 
Liposomes have already seen FDA approval for the delivery of small molecule 
delivery.89 Doxorubicin, an amphiphilic anticancer agent, has a great therapeutic 
potential, yet it is complicated by questionable low solubility, high toxicity 
and poor circulation. By loading in aggregates in the liposome core, it has 
been able to target tumors via the previously mentioned EPR effect with greater 
doses than previously possible. As illustrated by this example, the key advantage 
is the ability to enhance serum solubility of the small molecule drugs and 
achieve longer release profiles. In pulmonary settings, this has been used for the 
enhancement of free radical scavengers,90 enzyme inhibitors,91 and in anticancer 
treatments.92'93 
4.3. Types of nanocarriers 
Nanocarriers utilizing natural biomaterials or structures (e.g. liposomes consisting 
of natural phospholipids found in cellular plasma membranes) were the first to be 
explored for drug delivery.94 Since then, designs have included solid nanoparticles, 
double emulsion nanoparticles, polymeric micelles, polymersomes and worm-like 
micelles. Synthetic materials, especially polymeric materials, offer great freedom in 
that they can be designed to enhance circulation, reduce immunogenicity, provide 
environmentally responsive elements and possess biologically derived properties, 
also known as biomimetic properties, such as adhesion response elements and 
receptor ligands. All these carriers are amenable for pulmonary delivery. For a 
detailed review of the formation mechanisms and technical aspects of nanocarrier 
formulation, please refer to the reviews at Refs. 28, 95-98. 
4.4. Mechanisms of drug loading 
The main mechanisms for loading drugs into nanoparticles include surface absorption, 
aqueous inclusion, solid-phase immobilization, and complexation aggregates 
(Fig. 4). 
Surface absorption occurs via either hydrophobic interactions between the particle 
surface and hydrophobic interactions (e.g. tryptophan, tyrosine, phenylalanine 
for proteins) or electric charge interactions.99 This method is not effective for coating 
stealth nanocarriers, due to the nature of stealth mechanism, but can be used for 
the coupling of targeting moieties (see below) and therapeutic agents to non-stealth 
nanocarriers. 
In the context of pulmonary vascular targeting via IV route, stealth characteristics 
are not critically important due to the option of first pass delivery 
Delivery of Nanocarriers to the Lungs 513 
Surface Aqueous Solid-phase Complexation 
Absorption Inclusion Immobilization 
Fig. 4. Methods of nanocarriers loading with therapeutic agents. In the nano-scale range, 
surface absorption offers the greatest drug/particle loading, and most likely accounts for 
a fraction of loading in all reported nanocarriers, including those loaded by the alternative 
approaches. However, isolation of a cargo en route to target is most effective with inclusion 
mechanisms of loading. Currently, aqueous inclusion methods are most extensively explored 
for the loading of hydrophilic agents into polymer nanocarriers. Therapeutic effect may be 
achieved via either release of cargoes or diffusion of their substrates via polymer. Solid-phase 
immobilization is mostly used for loading of hydrophobic solutes, yet some proteins may 
also be amendable to this mechanism. Complexation relies upon the interaction of drug and 
polymer for particle formation, which permits formation of size-controled loaded vehicles. 
However, homogeneity of nanocarriers and drug release from these carriers are difficult to 
control. Carrier materials (e.g. polymers) are shown in a light grey color, drug loads are 
shown as dark spheres. 
mechanism. Indeed, latex poly(styrene) beads used as model prototype nonstealth 
nanocarriers (100 nm diameter) coated with surface-absorbed anti-ICAM, 
but not control IgG, showed very high pulmonary uptake after IV injection 
in mice.48 
Surface absorption does not protect a cargo from inactivation en route or in 
aggressive intracellular compartments (e.g. lysosomes), nor does it limit systemic 
side effects of circulating drugs. However, it may prolong circulation time, alter 
tissue targeting, and subsequently alter sub-cellular addressing of the drugs.15 It is 
the easiest method for nanocarrier loading with large MW drugs (e.g. therapeutic 
proteins).48,99-104 Latex beads surface coated with anti-ICAM and a therapeutic 
enzyme (catalase) provide a useful tool to the study of binding, internalization 
and degradation pathways for nanocarriers targeted to endothelial cells, the main 
cellular target in the pulmonary vasculature.48-102'105 
Liposomes can be loaded by aqueous core inclusion and by hydrophobic association 
within the lipid bilayer.19-106 Liposomes provide a large internal aqueous 
cargo compartment separated from milieu by the bilayer membrane. Since 
the cargo remains in an aqueous environment, its molecular mobility and enzymatic 
activity are not compromised. Liposomes afford effective loading and 
delivery of small hydrophobic agents (e.g. doxorubicin in Doxil®). In polymer 
nanocarriers, a polymer layer can provide even more protective barrier 
514 Dziubla & Muzykantov 
via either self-assembly mechanisms employed in synthesis of polymersomes,107 
double emulsion formation mechanism,88 or in nanoscale hydrogel synthesis 
techniques.108'109 
Solid-phase immobilization is an alternative strategy in which crystallized 
or lyophilized protein and small MW drugs are loaded as a suspension within 
the solid core of an organic, hydrophobic nanoparticle. High loadings of certain 
hydrophobic drugs have been reported. For instance, irinotecan, an anticancer 
therapeutic, was capable of being loaded at 4.5 wt% into 120 nm nanoparticles, 
composed of diblock PEG-poly(lactic-co-glycolic acid).110 This method may provide 
an added benefit in the delivery of bioactive drugs. The organic environment 
restricts mobility for some therapeutic protein resistant to unfolding, 
that may paradoxically yet simultaneously reduce activity and extend functional 
use.98,111,112 Moreover, since the protein is not in a soluble state, loading 
is not constrained by aqueous solubility limits and the entire particle core 
could support inclusion; hence, this mechanism may provide highly effective 
loading. 
The fourth mechanism for loading employs the complexation of a drug with the 
carrier material. Common approaches to complexation include inter-ionic associating 
mechanisms, the biotin-streptavidin cross-linking system, or covalent bonding. 
For instance, regular polymeric micelles of poly(ethylene glycol)-b-poly(aspartic 
acid) were formed in the presence of the positively charged lysozyme.113 Complexes 
can also take the form of polyplexes (e.g. poly(ethylimide) (PEI) and 
DNA), or in a single polymer chain coupling multiple proteins.64,114 This latter 
form has been popularized by the use of hydrophilic polymers such as poly(n(2- 
hydroxypropyl)methacrylamide) (HPMA), which uses amide linkages to covalently 
attach proteins and small molecules onto the polymer backbone.115 This also 
includes the polymer prodrugs that utilize degradable bonds to limit/control the 
therapeutic release rate.116,117 
Yet, even hydrophobic associations, disulfide linkages, streptavidin-biotin or 
antibody-antigen pairs can be used to form drug-polymer complexes. By controling 
the extent of modification of a therapeutic cargo and the affinity carrier by 
cross-linking agents and feed conditions, the complexation mechanism can result in 
nano-sized aggregates with a relatively high degree of drug inclusion.118 However, 
these conjugates (polyplexes) are characterized by significant heterogeneity, both in 
molecular composition and in size. Due to the nature of the conjugation mechanism, 
release from these systems is typically poor and mainly controled by degradation 
of the components. Thus, in the case of enzyme therapies (see Sec. 4.2.3), conjugates 
of this type, function effectively, typically only if enzymes substrates are small and 

diffusible enough to be accessible within the aggregate core, such as H2O2 in the 
case of catalase delivery.88 
Delivery of Nanocarriers to the Lungs 515 
4.5. Drug release mechanisms 
Nanocarriers can provide 3 main mechanisms of release for its drug cargo (Fig. 5). 
The most commonly considered release profile is that of continuous release (for a 
more detailed review, see Dziubla and Lowman119). Under this regimen, the drug 
slowly diffuses out of carrier particles over time, allowing for sustained high local 
concentrations of the drug. However, current nanocarrier formulations typically 
release ~ 40-70% of the total drug loaded within the first 6 hrs. This does not permit 
long-term therapy, but is rather suitable for therapies that require a burst release 
(e.g. gene and cancer treatment). 
Ideally, the cargo remains isolated from the systemic circulation and tissues 
until the intended target cells are reached and the release is triggered.120 This pattern 
allows for both the minimization of deleterious side effects, loss of activity, and 
(0 
ju 
tr 
3 
Q 
E 
O 
I 
E 
>« 
si 
c: 
IXI 
substrate 
Product 
Time 
Fig. 5. Modes of drug release. (A) Controled release allows for a therapeutic level of drug 
to be maintained for the greatest amount of time. (B) Delayed burst release is ideal for gene 
and cancer therapy, where immediate, local high concentrations are desired. (C) Sequestered 
enzyme delivery allows for a continuous activity of enzyme, even when the nanocarriers 
reside in compartments typically hostile to protein activity (e.g. lysosomes). In this scenario, 
carrier must be permeable for enzyme substrates or/and products. 
51 6 Dziubla & Muzykantov 
the minimization of the necessary effective dose. Finally, the nanocarrier may also 
be designed not to release the drug at all. For most instances, this prevents pharmacological 
activity. However, in the case of enzyme delivery where the substrate 
is diffusible (e.g. hydrogen peroxide, NO, oxygen, glucose, NAD) across the carrier 
wall, therapeutic activity may be achievable. This is especially suitable if the final 
targeting destination is lysosomes, which is likely to degrade the enzyme, thereby 
resulting in a loss of activity.88 
4.6. Nanocarriers for active targeting 
In order to achieve active targeting, affinity ligands are coupled to the surface 
of nanocarriers. Affinity and specificity of these ligands govern targeting. Yet, 
targeting of multivalent antibody-carrying nanoparticles differs from that of individual 
maternal antibodies in several important aspects. Firstly, high affinity of 
such complexes results in highly significant, in some instances, order of magnitude, 
enhancement of the pulmonary targeting of IV injected nanocarriers us maternal 
antibodies.44,47 Secondly, multivalent nanocarriers cross-link endothelial determinants, 
thus inducing highly effective endocytotic uptake, even though maternal 
antibodies are non internalizable.15'47'102,121 
Surface absorption, protein conjugation chemistries or biotin-streptavidin 
cross-linking can be utilized for the coupling of targeting entities, mainly monoclonal 
antibodies and their fragments to nanocarriers.9,105,122 Yet, the most important 
consideration is that of antibody presentation onto the carrier surface. For example, 
the antibodies attached covalently directly to the phospholipid head group of PEGylated 
liposomes, providing rather poor targeting due to the fact that extended PEG 
chains masked antibodies. This shortcoming can be solved by coupling the antibodies 
to the distal end of PEG chains. In fact, targeting of such stealth immunoliposomes 
exceeds that of standard liposomes, due to suppression of clearance 
mechanisms, and target group mobility and accessibility.19,122 
One of the most commonly employed conjugation strategies is that of 
maleimide sulfhydryl chemistry123 Maleimide group is more hydrolytically 
stable than other protein conjugation means, such as the amine directed 
n-hydroxysuccinate esters. Maleimide reacts with free thiol to create a non-reducible 
sulfide linkage. Since most proteins do not contain a free thiol group, competition 
between the drug (e.g. therapeutic protein) and the targeting moiety for available 
binding sites can be eliminated. 
Maleimide can be included onto the distal end of a PEG group in a PEG diblock 
copolymer.124,125 Upon nanoparticle synthesis, the PEG chain will extend out into 
the hydrophilic solution, ensuring the exposure of the maleimide group for subsequent 
conjugation. This allows for the separation of drug loading and nanocarrier 
Delivery of Nanocarriers to the Lungs 51 7 
formation from the conjugation of the targeting group. However, while maleimide 
hydrolysis is relatively slow at typical nanoparticle synthesis temperatures, it may 
still occur to a significant extent, thereby limiting the overall capacity for target 
group addition. 
5. Conclusion: Safety Issues, Limitations and Perspectives 
Results of in vitro and animal studies accumulated within the last decade strongly 
suggest that nanocarriers, especially those utilizing active targeting principles, will 
eventually provide a versatile and powerful technology platform for optimized 
drug delivery to the pulmonary vasculature. Extended surface of the pulmonary 
endothelium represents arguably the best target for drug delivery in the body, hence 
higher chances for sufficiently specific and effective drug delivery. 
On the other hand, in contrast with drug delivery to tumors, in which local 
toxic side effects can be considered as secondary benefits, safety of drug delivery 
to pulmonary vasculature is of greater concern. Thus, acute and delayed effects 
of targeting and endothelial uptake of nanocarriers on health and functions of the 
lung must be tested extremely rigorously. 
For example, pulmonary circulation is sensitive to subtle pro-inflammatory 
changes, often leading to edema and proliferation of sub-endothelial and interstitial 
components, pulmonary fibrosis and hypertension. In this context, an important 
question is how will nano-scale structures residing in a given pulmonary 
compartment, i.e. vascular lumen, lysosomes, be tolerated? How rapidly bloodstream 
and pulmonary lymphatic drainage can eliminate products of nanoparticles 
degradation? 
General safety concerns add to these specific issues that are pertinent to pulmonary 
targeting. Strictly speaking, the actual biocompatibility of materials for 
carriers remains unknown, until it is carefully tested in adequate clinical settings 
using carriers of adequate size. For example, nanocarriers based on polydactic 
glycolic acid) polymer, accepted for human use for macro-implants, may degrade 
into lactic acid and glycolic acid within the target cells, potentially exceeding its 
metabolic potential. Potentially harmful effects of activation systemic defense systems 
(i.e. complement, cytokines), overload of clearance systems (e.g. liver, kidneys) 
and immune reactions, represent general concerns of advanced delivery systems. 
However, despite these concerns, the most exciting prospect of nanocarriers are the 
near limitless possibilities for treatment strategies. Nanocarriers may be designed to 
contain multiple drugs, allowing for complex dosing regimes through just a single 
injection. 
Translational, industrial and commercial issues have to be addressed. For example, 
dosing (e.g. which drug load and particles dose afford therapeutic effects) and 
518 Dziubla & Muzykantov 
the timing of treatments have to be tested. Synthesis schemes and reagents readily 
adaptable to cGMP practices should be explored. Batch to batch variations and processing 
choices must be minimized, whereas the synthesis yield and drug loading 
effectiveness must be boosted to warrant practical utility. 
Targeting of nanocarriers to endothelial determinants in the pulmonary vasculature 
promises unprecedented levels of specificity and subcellular precision of 
drug delivery. Many endothelial determinants potentially useful for drug delivery 
including ecto-enzymes, cell adhesion molecules and caveolar antigens have 
been identified by methods including proteomics of endothelial plasma membrane, 
phage display libraries selections in vivo and the tracing of labeled antibodies. 
High-throughput, discovery-driven approaches such as phage display, map vascular 
lumen and identify novel targets enriched in defined areas of the lung or 
endothelial domains. Due to a limited insight into functions of these targets, some 
of them are unlikely to have a utility for drug delivery (e.g. due to safety concerns), 
yet all could be used as molecular probes in animal studies. 
Careful selection of targets and modulation of valency and size of the antibodydirected 
nanocarriers help to control intracellular uptake and traffic of cargoes. 
These parameters can be further fine-tuned, capitalizing on specific features of 
carriers including relatively labile protein conjugates, liposomes or polymer carriers 
with built-in rates of degradation and release, and membrane permeating moieties. 
It is tempting to speculate that the treatment of pathologies, including but not 
limited to acute lung injury, lung transplantation, pulmonary edema, thrombosis, 
hypertension and inflammation, will eventually benefit from targeting the delivery 
of drug nanocarriers to the pulmonary vasculature. 
Acknowledgments 
This work was supported by NHLBI ROl grants HL71175, HL078785 and HL73940, 
Department of Defense Grant (PR 012262) and Pennsylvania NTI core project. The 
authors thank Drs. S. Muro, M. Koval and V. Shuvaev (University of Pennsylvania) 
for the exciting and stimulating discussions and advice. 
References 
1. Muzykantov VR (2001) Delivery of antioxidant enzyme proteins to the lung. Antiox 
Redox Signal 3:39-62. 
2. Fishman AP (2004) A century of pulmonary hemodynamics. Am J Respir Crit Care Med 
170:109-113. 
3. Pandey R, Sharma A, Zahoor A, Sharma S, Khuller GK and Prasad B (2003) Poly (DLlactide-
co-glycolide) nanoparticle-based inhalable sustained drug delivery system for 
experimental tuberculosis. J Antimicrob Chemother 52:981-986. 
Delivery of Nanocarriers to the Lungs 519 
4. Corkery K (2000) Inhalable drugs for systemic therapy. Respir Care 45:831-835. 
5. Knecht A (1991) Inhalation therapy: alternative systems-an overview. / Aerosol Med 
4:189-192. 
6. Patton J (1998) Breathing life into protein drugs. Nat Biotechnol 16:141-143. 
7. Rau JL (2005) The inhalation of drugs: advantages and problems. Respir Care 50:367-382. 
8. Muzykantov VR (2003) Targeting pulmonary endothelium, in Muzykantov VR and 
Torchilin VP (eds.) Biomedical Aspects of Drug Targeting. Kluwer Academic Publishers: 
Boston, pp. 129-148. 
9. Muzykantov VR (2001) Targeting of superoxide dismutase and catalase to vascular 
endothelium. / Control Rel 71:1-21. 
10. Dziubla TD, Muro S, Muzykantov VR and Koval M (2005) Nanoscale Antioxidant 
Therapeutics, in Singh KK (ed.) Oxidative Stress, Disease and Cancer. Imperial College 
Press: London, in Press. 
11. Brantley-Sieders D, Parker M and Chen J (2004) Eph receptor tyrosine kinases in tumor 
and tumor microenvironment. Curr Pharm Des 10:3431-3442. 
12. Bibby DC, Talmadge JE, Dalai MK, Kurz SG, Chytil KM, Barry SE, Shand DG and Steiert 
M (2005) Pharmacokinetics and biodistribution of RGD-targeted doxorubicin-loaded 
nanoparticles in tumor-bearing mice, bit J Pharm 293:281-290. 
13. Stevens PJ, Sekido M and Lee RJ (2004) A folate receptor-targeted lipid nanoparticle 
formulation for a lipophilic paclitaxel prodrug. Pharm Res 21:2153-2157. 
14. Benns JM and Kim SW (2000) Tailoring new gene delivery designs for specific targets. 
/ Drug Targ 8:1-12. 
15. Muro S, Koval M and Muzykantov V (2004) Endothelial endocytic pathways: Gates for 
vascular drug delivery. Curr Vase Pharmacol 2:281-299. 
16. Murciano JC, Muro S, Koniaris L, Christofidou-Solomidou M, Harshaw DW, Albelda 
SM, Granger DN, Cines DB and Muzykantov VR (2003) ICAM-directed vascular 
immunotargeting of antithrombotic agents to the endothelial luminal surface. Blood 
101:3977-3984. 
17. Gref R, Minamitake Y, Peracchia MT, Trubetskoy V, Torchilin V and Langer R (1994) 
Biodegradable long-circulating polymeric nanospheres. Science 263:1600-1603. 
18. Kreuter J, Tauber U and Illi V (1979) Distribution and elimination of poly(methyl-2- 
14C-methacrylate) nanoparticle radioactivity after injection in rats and mice. / Pharm 
Sci 68:1443-1447. 
19. Moghimi SM, Hunter AC and Murray JC (2001) Long-circulating and target-specific 
nanoparticles: theory to practice. Pharmacol Rev 53:283-318. 
20. Harris JM (1992) Poly(ethylene Glycol) Chemistry: Biotechnical and Biomedical Applications. 
Plenum Press: New York. 
21. Moghimi SM and Szebeni J (2003) Stealth liposomes and long circulating nanoparticles: 
Critical issues in pharmacokinetics, opsonization and protein-binding properties. Prog 
Lipid Res 42:463^78. 
22. Forster S, Konrad M and Linder P (2005) Shear thinning and orientational ordering of 
wormlike micelles. Phys Rev Lett 94:017803. 
520 Dziubla & Muzykantov 
23. Li S, Nickels J and Palmer AF (2005) Liposome-encapsulated actin-hemoglobin 
(LEAcHb) artificial blood substitutes. Biomaterials 26:3759-3769. 
24. Li S and Palmer AF (2004) Structure of small actin-containing liposomes probed by 
atomic force microscopy: Effect of actin concentration & liposome size. Langmuir 
20:7917-7925. 
25. Yoon W (2004) Embolic agents used for bronchial artery embolisation in massive 
haemoptysis. Exp Opin Pharmacother 5:361-367. 
26. Kumar N, Ravikumar MN and Domb AJ (2001) Biodegradable block copolymers. Adv 
Drug De Rev 53:23-44. 
27. Heller J and Baker RW (1980) Theory and Practice of Controled Drug Delivery from Bioerodible 
Polymers, in Baker RW (ed.) Controled Release ofBioactive Materials. Academic 
Press: New York: 1-37. 
28. Dziubla TD and Muzykantov V (2006) Synthetic carriers for the delivery of protein 
therapeutics. Biotechnol Genet Eng Rev 22:267-299. 
29. Bragonzi A, Boletta A, Biffi A, Muggia A, Sersale G, Cheng SH, Bordignon C, Assael BM 
and Conese M (1999) Comparison between cationic polymers and lipids in mediating 
systemic gene delivery to the lungs. Gene Titer 6:1995-2004. 
30. Floch V, Delepine P, Guillaume C, Loisel S, Chasse S, Le Bolc'h G, Gobin E, Leroy JP and 
Ferec C (2000) Systemic administration of cationic phosphonolipids/DNA complexes 
and the relationship between formulation and lung transfection efficiency. Biochim 
Biophys Acta 1464:95-103. 
31. Ravi Kumar MN, Sameti M, Mohapatra SS, Kong X, Lockey RF, Bakowsky U, 
Lindenblatt G, Schmidt H and Lehr CM (2004) Cationic silica nanoparticles as gene 
carriers: Synthesis, characterization and transfection efficiency in vitro and in vivo. } 
Nanosci Nanotechnol 4:876-881. 
32. Kwok KY, Yang Y and Rice KG (2001) Evolution of cross-linked non-viral gene delivery 
systems. Curr Opin Mol Ther 3:142-146. 
33. Benigni A, Tomasoni S and Remuzzi G (2002) Impediments to successful gene transfer 
to the kidney in the context of transplantation and how to overcome them. Kidney lnt 
61:115-119. 
34. Thierry AR, Lunardi-Iskandar Y, Bryant JL, Rabinovich P, Gallo RC and Mahan LC 
(1995) Systemic gene therapy: Biodistribution and long-term expression of a transgene 
in mice. Proc Natl Acad Sci USA 92:9742-9746. 
35. Rodman DM, San H, Simari R, Stephan D, Tanner F, Yang Z, Nabel GJ and Nabel EG 
(1997) In vivo gene delivery to the pulmonary circulation in rats: transgene distribution 
and vascular inflammatory response. Am J Respir Cell Mol Biol 16:640-649. 
36. Torchilin VP (2000) Drug targeting. Eur ] Pharm Sci ll(Suppl 2):S81-S91. 
37. Fang J, Sawa T and Maeda H (2003) Factors and mechanism of "EPR" effect and the 
enhanced antitumor effects of macromolecular drugs including SMANCS. Adv Exp Med 
Biol 519:29-49. 
38. Maeda H, Fang J, Inutsuka T and Kitamoto Y (2003) Vascular permeability enhancement 
in solid tumor: Various factors, mechanisms involved and its implications. Int 
Immunopharmacol 3:319-328. 
Delivery of Nanocarriers to the Lungs 521 
39. Oh P, Li Y, Yu J, Durr E, Krasinska KM, Carver LA, Testa JE and Schnitzer JE (2004) 
Subtractive proteomic mapping of the endothelial surface in lung and solid tumours 
for tissue-specific therapy. Nature 429:629-635. 
40. Muzykantov VR, Puchnina EA, Atochina EN, Hiemish H, Slinkin MA, Meertsuk FE 
and Danilov SM (1991) Endotoxin reduces specific pulmonary uptake of radiolabeled 
monoclonal antibody to angiotensin-converting enzyme. / Nucl Med 32:453-460. 
41. Trubetskoy VS, Torchilin VP, Kennel SJ and Huang L (1992) Use of N-terminal modified 
poly(L-lysine)-antibody conjugate as a carrier for targeted gene delivery in mouse lung 
endothelial cells. Bioconjug Chem 3:323-327. 
42. Schuchman E and Muro S (2005) The development of enzyme replacement therapy for 
lysosomal diseases: Gaucher disease and beyond, in Futerman T and Zimran A (eds.) 
Gaucher Disease: Lessons Learned About Therapy of Lysosomal Diseases. CRC Press: Boca 
Raton, in Press. 
43. Rajotte D, Arap W, Hagedorn M, Koivunen E, Pasqualini R and Ruoslahti E (1998) 
Molecular heterogeneity of the vascular endothelium revealed by in vivo phage display. 
/ Clin Invest 102:430-437. 
44. Danilov SM, Gavrilyuk VD, Franke FE, Pauls K, Harshaw DW, McDonald TD, Miletich 
DJ and Muzykantov VR (2001) Lung uptake of antibodies to endothelial antigens: Key 
determinants of vascular immunotargeting. Am J Physiol 280:L1335-L1347. 
45. Danilov S, Atochina E, Hiemisch H, Churak-ova T, Moldobayeva A, Sakharov I, 
Deichman G, Ryan U and Muzykantov VR (1994) Interaction of mAb to angiotensinconverting 
enzyme (ACE) with antigen in vitro and in vivo: Antibody targeting to the 
lung induces ACE antigenic modulation. Int Immunol 6:1153-11560. 
46. Danilov SM, Muzykantov VR, Martynov AV, Atochina EN, Sakharov IY, Trakht IN and 
Smirnov VN (1991) Lung is the target organ for a monoclonal antibody to angiotensinconverting 
enzyme. Lab Invest 64:118-124. 
47. Muzykantov VR, Christofidou-Solomidou M, Balyasnikova I, Harshaw DW, Schultz 
L, Fisher AB and Albelda SM (1999) Streptavidin facilitates internalization and pulmonary 
targeting of an anti-endothelial cell antibody (platelet-endothelial cell adhesion 
molecule (1): A strategy for vascular immunotargeting of drugs. Proc Nat Acad Sci 
USA 96:2379-2384. 
48. Muro S, Gajewski C, Koval M and Muzykantov VR (2005) ICAM-1 recycling in endothelial 
cells: A novel pathway for sustained intracellular delivery and prolonged effects of 
drugs. Blood 105:650-658. 
49. Kelly KA, AUport JR, Tsourkas A, Shinde-Patil VR, Josephson L and Weissleder R 
(2005) Detection of vascular adhesion molecule-1 expression using a novel multimodal 
nanoparticle. Circ Res 96:327-336. 
50. Sakhalkar HS, Dalai MK, Salem AK, Ansari R, Fu J, Kiani MF, Kurjiaka DT, Hanes J, 
Shakesheff KM and Goetz DJ (2003) Leukocyte-inspired biodegradable particles that 
selectively and avidly adhere to inflamed endothelium in vitro and in vivo. Proc Natl 
Acad Sci USA 100:15895-15900. 
51. Ogawara K, Rots MG, Kok RJ, Moorlag HE, Van Loenen AM, Meijer DK, Haisma HJ 
and Molema G (2004) A novel strategy to modify adenovirus tropism and enhance 
522 Dziubla & Muzykantov 
transgene delivery to activated vascular endothelial cells in vitro and in vivo. Hum Gene 
Ther 15:433-443. 
52. Spragg DD, Alford DR, Greferath R, Larsen CE, Lee KD, Gurtner GC, Cybulsky MI, Tosi 
PF, Nicolau C and Gimbrone MA, Jr. (1997) Immunotargeting of liposomes to activated 
vascular endothelial cells: A strategy for site-selective delivery in the cardiovascular 
system. Proc Natl Acad Sci USA 94:8795-8800. 
53. Lindner JR, Song J, Christiansen J, Klibanov AL, Xu F and Ley K (2001) Ultrasound 
assessment of inflammation and renal tissue injury with microbubbles targeted to 
P-selectin. Circulation 104:2107-2112. 
54. Predescu D, Predescu S and Malik AB (2002) Transport of nitrated albumin across 
continuous vascular endothelium. Proc Natl Acad Sci USA 99:13932-13937. 
55. Durr E, Yu J, Krasinska KM, Carver LA, Yates JR, Testa JE, Oh P and Schnitzer JE (2004) 
Direct proteomic mapping of the lung microvascular endothelial cell surface in vivo 
and in cell culture. Nat Biotechnol 22:985-992. 
56. Mcintosh DP, Tan XY, Oh P and Schnitzer JE (2002) Targeting endothelium and its 
dynamic caveolae for tissue-specific transcytosis in vivo: A pathway to overcome cell 
barriers to drug and gene delivery. Proc Natl Acad Sci USA 99:1996-2001. 
57. Disegi JA (2000) Titanium alloys for fracture fixation implants. Injury 31 (Suppl 4):14-17. 
58. Gotman I (1997) Characteristics of metals used in implants. / Endourol 11: 
383-389. 
59. Gole J, Burda C, Fedorov A and White M (2003) Enhanced reactivity and phase transformation 
at the nanoscale: Efficient formation of active silica and doped and metal 
seeded Ti02-xNx photocatalysts. Rev Adv Mater Sci 5:265-269. 
60. Oberdorster G, Finkelstein JN, Johnston C, Gelein R, Cox C, Baggs R and Elder AC 
(2000) Acute pulmonary effects of ultrafine particles in rats and mice. Res Rep Health 
Eff Inst 5-74; 75-86. 
61. Warner WG, Yin JJ and Wei RR (1997) Oxidative damage to nucleic acids photosensitized 
by titanium dioxide. Free Radic Biol Med 23:851-858. 
62. Borm PJ and Kreyling W (2004) Toxicological hazards of inhaled nanoparticles — 
potential implications for drug delivery. / Nanosci Nanotechnol 4:521-531. 
63. Thanou M and Duncan R (2003) Polymer-protein and polymer-drug conjugates in 
cancer therapy. Curr Opin Invest Drugs 4:701-709. 
64. Duncan R (2003) The dawning era of polymer therapeutics. Nat Rev Drug Discov 2:347- 
360. 
65. Muller NL (2002) Computed tomography and magnetic resonance imaging: past, 
present and future. Eur Respir } (Suppl)35:3s-12s. 
66. Kauczor HU and Kreitner KF (2000) Contrast-enhanced MRI of the lung. Eur } Radiol 
34:196-207. 
67. Mclntire GL, Bacon ER, Illig KJ, Coffey SB, Singh B, Bessin G, Shore MT and Wolf GL 
(2000) Time course of nodal enhancement with CT X-ray nanoparticle contrast agents: 
Effect of particle size and chemical structure. Invest Radiol 35:91-96. 
68. Wisner ER, Katzberg RW, Koblik PD, Shelton DK, Fisher PE, Griffey SM, Drake C, 
Harnish PP, Vessey AR, Haley PJ, et al. (1994) Iodinated nanoparticles for indirect 
Delivery of Nanocarriers to the Lungs 523 
computed tomography lymphography of the craniocervical and thoracic lymph nodes 
in normal dogs. Acad Radiol 1:377-384. 
69. Spuentrup E, Buecker A, Katoh M, Wiethoff AJ, Parsons EC, Jr., Botnar RM, Weisskoff 
RM, Graham PB, Manning WJ and Gunther RW (2005) Molecular magnetic resonance 
imaging of coronary thrombosis and pulmonary emboli with a novel fibrin-targeted 
contrast agent. Circulation 111:1377-1382. 
70. Choi H, Choi SR, Zhou R, Kung HF and Chen IW (2004) Iron oxide nanoparticles 
as magnetic resonance contrast agent for tumor imaging via folate receptor-targeted 
delivery. Acad Radiol 11:996-1004. 
71. Martins S, de Perrot M, Imai Y, Yamane M, Quadri SM, Segall L, Dutly A, Sakiyama S, 
Chaparro A, Davidson BL, Waddell TK, Liu M and Keshavjee S (2004) Transbronchial 
administration of adenoviral-mediated interleukin-10 gene to the donor improves function 
in a pig lung transplant model. Gene They 11:1786-1796. 
72. Suda T, Mora BN, D'Ovidio F, Cooper JA, Hiratsuka M, Zhang W, Mohanakumar T 
and Patterson GA (2000) In vivo adenovirus-mediated endothelial nitric oxide synthase 
gene transfer ameliorates lung allograft ischemia-reperfusion injury / Thorac Cardiovasc 
Surg 119:297-304. 
73. Hiratsuka M, Mora BN, Yano M, Mohanakumar T and Patterson GA (1999) Gene transfer 
of heat shock protein 70 protects lung grafts from ischemia-reperfusion injury. Ann 
Thorac Surg 67:1421-1427. 
74. Reynolds PN, Nicklin SA, Kaliberova L, Boatman BG, Grizzle WE, Balyasnikova IV, 
Baker AH, Danilov SM and Curiel DT (2001) Combined transductional and transcriptional 
targeting improves the specificity of transgene expression in vivo. Nat Biotechnol 
19:838-842. 
75. Reynolds PN, Zinn KR, Gavrilyuk VD, Balyasnikova IV, Rogers BE, Buchsbaum DJ, 
Wang MH, Miletich DJ, Grizzle WE, Douglas JT, Danilov SM and Curiel DT (2000) 
A targetable, injectable adenoviral vector for selective gene delivery to pulmonary 
endothelium in vivo. Mol Ther 2:562-578. 
76. Wagner E (2004) Strategies to improve DNA polyplexes for in vivo gene transfer: Will 
"artificial viruses" be the answer? Pharm Res 21:8-14. 
77. Elouahabi A and Ruysschaert JM (2005) Formation and intracellular trafficking of 
lipoplexes and polyplexes. Mol Ther 11:336-347. 
78. Rodriguez EG (2004) Nonviral DNA vectors for immunization and therapy: Design 
and methods for their obtention. / Mol Med 82:500-509. 
79. Griesenbach U, Geddes DM and Alton EW (2004) Advances in cystic fibrosis gene 
therapy. Curr Opin Pulm Med 10:542-546. 
80. Barron LG, Uyechi LS and Szoka FC, Jr. (1999) Cationic lipids are essential for gene 
delivery mediated by intravenous administration of lipoplexes. Gene Ther 6:1179-1183. 
81. Li S, Tan Y, Viroonchatapan E, Pitt BR and Huang L (2000) Targeted gene delivery to 
pulmonary endothelium by anti-PECAM antibody. Am J Physiol Lung Cell Mol Physiol 
278:L504-511. 
82. Balyasnikova IV, Yeomans DC, McDonald TB and Danilov SM (2002) Antibodymediated 
lung endothelium targeting: In vivo model on primates. Gene Ther 9:282-290. 
524 Dziubla & Muzykantov 
83. Balyasnikova IV, Metzger R, Visintine DJ, Dimasius V, Sun ZL, Berestetskaya YV, 
McDonald TD, Curiel DT, Minshall RD and Danilov SM (2005) Selective rat lung 
endothelial targeting with a new set of monoclonal antibodies to angiotensin 
I-converting enzyme. Pulm Pharmacol Ther 18:251-267. 
84. Gopalan B, Ito I, Branch CD, Stephens C, Roth JA and Ramesh R (2004) Nanoparticle 
based systemic gene therapy for lung cancer: Molecular mechanisms and strategies 
to suppress nanoparticle-mediated inflammatory response. Technol Cancer Res Treat 3: 
647-657. 
85. Li S, Wu SP, Whitmore M, Loeffert EJ, Wang L, Watkins SC, Pitt BR and Huang L (1999) 
Effect of immune response on gene transfer to the lung via systemic administration of 
cationic lipidic vectors. Am J Physiol 276:L796-804. 
86. Kozower BD, Christofidou-Solomidou M, Sweitzer TD, Muro S, Buerk DG, Solomides 
CC, Albelda SM, Patterson GA and Muzykantov VR (2003) Immunotargeting of catalase 
to the pulmonary endothelium alleviates oxidative stress and reduces acute lung 
transplantation injury. Nat Biotechnol 21:392-398. 
87. Christofidou-Solomidou M, Scherpereel A, Wiewrodt R, Ng K, Sweitzer T, Arguiri E, 
Shuvaev V, Solomides CC, Albelda SM and Muzykantov VR (2003) PECAM-directed 
delivery of catalase to endothelium protects against pulmonary vascular oxidative 
stress. Am } Physiol 285:L283-L292. 
88. Dziubla TD, Karim A and Muzykantov VR (2005) Polymer nanocarriers protecting 
active enzyme cargo against proteolysis. / Control Rel 102:427-439. 
89. Abraham SA, Waterhouse DN, Mayer LD, Cullis PR, Madden TD and Bally MB (2005) 
The liposomal formulation of doxorubicin. Meth Enzymol 391:71-97. 
90. Stone WL and Smith M (2004) Therapeutic uses of antioxidant liposomes. Mol Biotechnol 
27:217-230. 
91. Spina D (2003) Phosphodiesterase-4 inhibitors in the treatment of inflammatory lung 
disease. Drugs 63:2575-2594. 
92. Peer D and Margalit R (2004) Tumor-targeted hyaluronan nanoliposomes increase the 
antitumor activity of liposomal Doxorubicin in syngeneic and human xenograft mouse 
tumor models. Neoplasia 6:343-353. 
93. Lu B, Zhang JQ and Yang H (2003) Lung-targeting microspheres of carboplatin. bit J 
Pharm 265:1-11. 
94. Mainardes RM and Silva LP (2004) Drug delivery systems: Past, present, and future. 
Curr Drug Targ 5:449^55. 
95. Panyam J and Labhasetwar V (2003) Biodegradable nanoparticles for drug and gene 
delivery to cells and tissue. Adv Drug Del Rev 55:329-347. 
96. Ulrich AS (2002) Biophysical aspects of using liposomes as delivery vehicles. Biosci Rep 
22:129-150. 
97. Hans ML and Lowman AM (2002) Biodegradable nanoparticles for drug delivery and 
targeting. Current Opin Solid State Mater Sci 6:319-327. 
98. Discher DE and Eisenberg A (2002) Polymer Vesicles. Science 297:967-973. 
Delivery of Nanocarriers to the Lungs 525 
99. Sakuma S, Suzuki N, Sudo R, Hiwatari K, Kishida A and Akashi M (2002) Optimized 
chemical structure of nanoparticles as carriers for oral delivery of salmon calcitonin. 
IntJPharm 239:185-195. 
100. Constancis A, Meyrueix R, Bryson N, Huille S, Grosselin JM, Gulik-Krzywicki T and 
Soula G (1999) Macromolecular colloids of diblock poly(amino acids) that bind insulin. 
/ Coll Interf Sci 217:357-368. 
101. Lvov Y and Caruso F (2001) Biocolloids with ordered urease multilayer shells as enzymatic 
reactors. Anal Chem 73:4212^217. 
102. Muro S, Wiewrodt R, Thomas A, Koniaris L, Albelda SM, Muzykantov VR and 
Koval M (2003) A novel endocytic pathway induced by clustering endothelial ICAM-1 
or PECAM-1. / Cell Sci 116:1599-1609. 
103. Michaelis M, Matousek J, Vogel JU, Slavik T, Langer K, Cinatl J, Kreuter J and Schwabe D 
(2000) Bovine seminal ribonuclease attached to nanoparticles made of polylactic acid 
kills leukemia and lymphoma cell lines in vitro. Anticancer Drugs 11:369-376. 
104. Bousquet Y, Swart PJ, Schmitt-Colin N, Velge-Roussel F, Kuipers ME, Meijer DK, Bru N, 
Hoebeke J and Breton P (1999) Molecular mechanisms of the adsorption of a model 
protein (human serum albumin) on poly(methylidene malonate 2.1.2) nanoparticles. 
Pharm Res 16:141-147. 
105. Muro S, Muzykantov VR and Murciano JC (2004) Characterization of endothelial internalization 
and targeting of antibody-enzyme conjugates in cell cultures and in laboratory 
animals. Meth Mol Biol 283:21-36. 
106. Ceh B, Winterhalter M, Frederik PM, Vallner JJ and Lasic DD (1997) Stealth Liposomes: 
From theory to product. Adv Drug Del Rev 24:165-177. 
107. Discher BM, Won Y-Y, Ege DS, Lee CC-M, Bates FS, Discher DE and Hammer DA (1999) 
Polymersomes: Tough vesicles made from Diblock copolymers. Science 284:143-146. 
108. Huang G, Gao J, Hu Z, St John JV, Ponder BC and Moro D (2004) Controlled drug 
release from hydrogel nanoparticle networks. / Control Rel 94:303-311. 
109. Peppas NA, Wood KM and Blanchette JO (2004) Hydrogels for oral delivery of therapeutic 
proteins. Exp Opin Biol Ther 4:881-887. 
110. Onishi H and Machida Y (2003) Antitumor properties of irinotecan-containing 
nanoparticles prepared using poly(DL-lactic acid) and poly(ethylene glycol)-blockpoly(
propylene glycol)-block-poly(ethylene glycol). Biol Pharm Bull 26:116-119. 
111. Klibanov AM (2001) Improving enzymes by using them in organic solvents. Nature 
409:241-246. 
112. Klibanov AM (1997) Why are enzymes less active in organic solvents than in water? 
Trends Biotechnol 15:97-101. 
113. Harada A and Kataoka K (2001) Pronounced activity of enzymes through the incorporation 
into the core of polyion complex micelles made from charged block copolymers. 
J Control Rel 72:85-91. 
114. Godbey WT, Wu KK and Mikos AG (1999) Poly(ethylenimine) and its role in gene 
delivery. / Control Rel 60:149-160. 
526 Dziubla & Muzykantov 
115. Kopecek J, Kopeckova P, Minko T, Lu ZR and Peterson CM (2001) Water soluble polymers 
in tumor targeted delivery. / Control Rel 74:147-158. 
116. Hoste K, De Winne K and Schacht E (2004) Polymeric prodrugs. IntJPharm 277:119-131. 
117. Ulbrich K and Subr V (2004) Polymeric anticancer drugs with pH-controlled activation. 
Adv Drug Del Rev 56:1023-1050. 
118. Shuvaev VV, Dziubla T, Wiewrodt R and Muzykantov VR (2004) Streptavidin-biotin 
crosslinking of therapeutic enzymes with carrier antibodies: Nanoconjugates for protection 
against endothelial oxidative stress. Meth Mol Biol 283:3-19. 
119. Dziubla TD and Lowman AM (2001) Gels, in Schwartz M (ed.) Encyclopedia of Smart 
Materials. Wiley and Sons: New York: 1-12. 
120. Lowman AM, Dziubla TD, Bures P and Peppas NA (2004) Structural and dynamic 
response of neutral and intelligent networks in biomedical environments, in Peppas 
NA and Sefton MV (eds.) Advances In Chemical Engineering: Molecular and Cellular Foundations 
of Biomaterials. Academic Press: New York, 29:75-122. 
121. Wiewrodt R, Thomas AP, Cipelletti L, Christofidou-Solomidou M, Weitz DA, 
Feinstein SI, Schaffer D, Albelda SM, Koval M and Muzykantov VR (2002) Sizedependent 
intracellular immunotargeting of therapeutic cargoes into endothelial cells. 
Blood 99:912-922. 
122. Torchilin VP (1994) Immunoliposomes and PEGylated immunoliposomes: Possible use 
for targeted delivery of imaging agents. Immunomethods 4:244-258. 
123. Hermanson GT (1996) Bioconjugate Techniques. Acedemic Press: San Diego, CA. 
124. Olivier JC, Huertas R, Lee HJ, Calon F and Pardridge WM (2002) Synthesis of pegylated 
immunonanoparticles. Pharm Res 19:1137-1143. 
125. Tessmar J, Mikos A and Gopferich A (2003) The use of poly(ethylene glycol)-blockpoly(
lactic acid) derived copolymers for the rapid creation of biomimetic surfaces. 
Biomaterials 24:4475-4486. 
24 
Nanoparticulate Carriers for Drug 
Delivery to the Brain 
Jorg Kreuter 
1. Introduction 
The following chapter deals with a subject that according to the journals Science or 
Nature, is of no general interest, namely the brain or to be more specific, drug delivery 
to the brain (personal communication). The brain is one of the best protected 
organs of the body, to the outside by the skull and towards the blood circulation by 
the blood-brain barrier (BBB). The purpose of the BBB is to maintain the homeostasis 
of the brain, and to allow the creation of a unique extracellular fluid environment 
within the central nervous system (CNS), whose composition can as a consequence 
be precisely controlled.1 The extracellular fluid compartments of the CNS comprise 
of the brain, the spinal cord parenchymal interstitial fluid and the cerebrospinal 
fluid contained within the ventricles of the brain, as well as the cerebral and spinal 
subarachnoid spaces. The structural BBB is created by the endothelial cells forming 
the capillaries of the brain and the spinal cord.1 These endothelial cells are characterized 
by having tight continuous circumferential junctions between them, thus 
abolishing any aqueous paracellular pathways between these cells.2 The presence of 
the tight junctions and the lack of aqueous pathways between cells greatly restricts 
the movement of polar solutes across the cerebral endothelium.3 Certain substances 
may diffuse passively across the brain endothelial cells. This diffusion is dependent 
on lipophilicity and molecular weight. Drugs with a molecular weight above 500 Da 
are normally excluded from a passive diffusional transport across the BBB. 
527 
528 Kreuter 
However, a large number of drugs that would possess a favorable lipophilicity 
and molecular weight, which should normally enable an easy transport across 
the BBB, are rapidly pumped back into the blood stream by extremely effective 
efflux pumps.3 - 5 These pump systems include among others, P-glycoprotein 
(Pgp), also referred to as multidrug resistance protein (MDR), as well as MOAT 
(multiple organic anion transporter). Since the brain is dependent on the blood 
to deliver substrates as well as to remove metabolic waste, the endothelial cells 
are also required to maintain a high level of carrier-mediated transport systems 
that enable the entry or the elimination of a variety of endogenous compounds 
including hydrophilic substances such as hexoses, amino acids, purine 
compounds, and mono-carbonic substances,6 as well as lipoproteins including LDL 
(low density lipoprotein).7,8 Some of these transporters are unidirectional and some 
bi-directional in their transport of solutes across the cell membrane. This polarization 
means that some solutes can be preferentially transported into or out of 
the brain.1 
As a consequence, the BBB presents a huge challenge for the effective delivery 
of a large number of therapeutics to the brain, and, therefore, many attempts 
have been made to overcome this barrier. For instance, these attempts include the 
osmotic opening of the tight junctions,9,10 use of prodrugs or carrier systems such as 
antibodies,11,12 liposomes13,14 and nanoparticles. Opening of the tight junctions by 
osmotic pressure, however, is a very invasive procedure that also enables the entry 
of unwanted substances into the brain. The employment of prodrugs may yield 
a higher lipophilicity, enabling a better permeation and transport into and across 
the lipophilic endothelial barrier, and/or these prodrugs may use the membrane 
associated carrier systems. In many cases, however, a suitable prodrug cannot be 
synthesized, or the resulting molecular weight is too large. Colloidal carriers also 
can take advantage of these carrier systems present in the BBB. These systems, for 
instance, include the lipoprotein receptors and the transferrin transcytosis systems, 
and may be employed in the delivery of drugs by the above particulate colloidal 
drug delivery systems. 
2. Nanoparticles 
Nanoparticles for pharmaceutical purposes as defined by the Encyclopedia of Pharmaceutical 
Technology15 are solid colloidal particles ranging in size from 1 to 
1000 nm (1 /xm), consisting of macromolecular materials in which the active principle 
(drug or biologically active material) is dissolved, entrapped, or encapsulated, 
or to which the active principle is adsorbed or attached. The use of nanoparticles for 
the transport of drugs across the BBB was already suggested in the early 1980s by 
Prof Speiser at the ETH (Swiss Federal Institute of Technology) in Zurich (personal 
Nanoparticulate Carriers for Drug Delivery to the Brain 529 
communication), who was also the first to systematically develop nanoparticles for 
drug delivery purposes in the late 1960s and early 1970s.15-17 
The possibility to use nanoparticles for the transport of drugs into the brain 
across the BBB was first shown with the hexapeptide dalargin (Tyr-D-Ala-Gly-Phe- 
Leu-Arg), a Leu-enkephalin analogue with opioid activity.18'19 This drug was bound 
to nanoparticles of a size of about 250 nm, made of the very rapidly biodegradable 
polymer poly (butyl cyanoacrylate). This material is one of the most rapidly 
biodegradable nanoparticle materials.20 The nanoparticles were incubated with this 
drug for 4 hrs, yielding the sorptive binding of 40% of the initial amount of dalargin. 
Overcoating of these particles with the surfactant polysorbate 80 (Tween® 80) was 
then achieved by further incubation for another 30 min with this surfactant, resulting 
in an equilibrium between surface-bound polysorbate and polysorbate in solution. 
A dose-dependent antinociceptive (analgesic) effect was observed using the 
tail-flick test, after intravenous injection to mice (Fig. 1) which was later repeated 
by other research groups with the hot plate test.21,22 The antinociceptive effect was 
accompanied by a pronounced Straub effect and could be totally inhibited by injection 
of the ^-opiate receptor antagonist naloxone 10 min before the injection of 
the nanoparticle preparation. Both results indicate a central action of dalargin on 
the CNS, demonstrating that it was indeed transported across the BBB and that 
the observed antinociceptive effects were not due to peripheral activity. In contrast 
to the polysorbate 80-coated nanoparticles, none of the controls including 
LU 
Q. 
2 
Dal saline Dal + Ps80 Dal-NP 
(10mg/kg) (10mg/kg) (10mg/kg) 
Dal-NP Dal-NP 
+Ps80 +Ps80 
(2.5 mg/kg) (5 mg/kg) 
Dal-NP 
+Ps80 
(7.5 mg/kg) 
Fig. 1. Antinociceptive effects after intravenous injection of different dalargin (Dal) formulations 
into mice. MPE = maximal possible effect; NP = nanoparticles; Ps 80 = polysorbate 80. 
530 Kreuter 
a solution of dalargin, a solution of polysorbate 80, a suspension of poly (butyl 
cyanoacrylate) nanoparticles, a mixture of dalargin with polysorbate 80, dalargin 
with nanoparticles or a mixture of all three components, dalargin, polysorbate 80, 
and nanoparticles, mixed immediately before injection, as well as dalargin bound to 
nanoparticles without polysorbate 80 coating, were able to exhibit any antinociceptive 
action (Fig. 1). The antinociceptive effects also showed circadian phase-(daytime)-
dependency, as well as a shift of the minima and maxima of the nociceptive 
reactions of the mice of almost 12 hrs compared with the controls and the dalargin 
solution.22 
3. Biodistribution 
3.1. Influence of surfactants on the biodistribution 
of nanoparticles 
Fundamental biodistribution studies of Troster et al.23 with 14C-labelled 
poly(methyl methacrylate) nanoparticles demonstrated that the coating of these 
nanoparticles with certain surfactants increased the whole brain concentrations 
of the nanoparticles in rats after intravenous injection. However, at that time, the 
authors were convinced that the nanoparticles were not taken up by any brainassociated 
cells, including the brain capillary endothelial cells, nor were transported 
across the BBB, but rather remained in the blood lumen adhering to the luminal 
surface of the endothelial cells.23 In addition, it has to be noted that some surfactants 
in Troster's experiments led to high [14C] brain concentrations, which were unable 
to achieve any antinociceptive effects with dalargin bound to the nanoparticles.24 
These important antinociceptive effects in the CNS with the polysorbate 80- 
coated dalargin nanoparticles reported above (Sec. 24.2) led to the investigation 
of the biodistribution of this drug after intravenous injection to mice, using 3Hlabelled 
dalargin in the form of [Leucyl-4,5-3H]-dalargin25 as well as of [3H-Tyr]- 
dalargin.26 Up to 3-fold higher concentrations in brain homogenates were found 
with the polysorbate 80-coated nanoparticles than with dalargin solution. These 
concentration differences were statistically different at most time points, although 
smaller than expected from the huge difference in the pharmacological responses. 
However, it has to be considered that the determination of the 3H-radioactivity in 
brain homogenates cannot discriminate between drug that has and drug that has 
not actually crossed the BBB. In addition, the observed concentration differences 
between different brain homogenate fractions25 may be the reason for the lack of 
efficient BBB crossing of dalargin in solution form. 
Much larger and important brain concentration differences were obtained 
after intravenous injection of poly(butyl cyanoacrylate) nanoparticles loaded with 
doxorubicin.27 In this case, the drug was added during polymerization. Four 
Nanoparticulate Carriers for Drug Delivery to the Brain 531 
doxorubicin formulations, (1) doxorubicin solution in saline, (2) doxorubicin solution 
plus 1% polysorbate 80 in saline, (3) doxorubicin bound to nanoparticles, or 
(4) doxorubicin bound to nanoparticles coated with polysorbate 80, were injected 
into the tail vein of rats at a doxorubicin dosage of 5 mg/kg. In the brain, the polysorbate 
80-coated nanoparticles yielded high doxorubicin concentrations of 6 /xg/g 
tissue between 2 and 4 hrs after injection. The brain concentrations were still at a 
level of about 1 /xg/g after 8 hrs, while the three other preparations remained below 
the detection limit of 0.1 /xg/g at all times. In contrast, very low concentration differences 
appeared between all four preparations in the blood only. Interestingly, the 
heart concentrations of both nanoparticle formulations remained very low, confirming 
earlier results of Couvreur et al.,2& whereas the heart concentrations with 
the two solutions were about 17 times higher than with the nanoparticles. Since the 
use of doxorubicin is limited by its cumulative high heart toxicity, this observation 
is of major significance. 
Solid lipid nanoparticles (SLN) were also able to achieve significant brain concentrations 
after intravenous, and even after duodenal administration. SLNs consisting 
of stearic acid, the surfactant Epicuron® 200, and taurocholate sodium loaded 
with doxorubicin,29 tobramycin,30 or idarubicin31 were prepared by dispersing a 
microemulsion containing the above components in water. At a dose of 6 mg/kg 
doxorubicin, brain concentrations of about 2 Mg/g were obtained after 180 min only 
with the SLNs, and no doxorubin was detectable in the brain of rats after i.v. administration 
of the solution through the jugular vein.29 With tobramycin (5 mg/kg), the 
intravenous route was compared with duodenal administration through a surgically 
implanted cannula. No tobramycin was detectable in the brain after administration 
of tobramycin solution to the rats. However, with the solid lipid nanoparticles, the 
amount of tobramycin in the brain 4 hrs after duodenal administration (4.8 /xg/g) 
was comparable to that after i.v. administration (4.5/xg/g). The tobramycin brain 
concentration was decreased 24 hrs after duodenal dosing, while the levels after i.v. 
administration remained fairly high (5.1 Mg/g)- In all other tissues except the brain, 
the tobramycin levels were higher after i.v. administration of the solution than those 
obtained with the solid lipid nanoparticles.30 Duodenal administration of idarubicin, 
at a dose of lmg/kg bound to solid lipid nanoparticles, yielded brain concentrations 
of about 11.2 n g /g after 24 hrs. This concentration was similar to that in the 
heart (11.5 ng/g) and about half of that in the liver. No detectable idarubicin nor the 
metabolit idarubicinol was found in the brain after administration of the solution.31 
Solid lipid nanoparticles consisting of stearic acid, soybean lecithin, and the 
surfactant poloxamer 188 (Pluronic® F68) loaded with the anticancer drug camptothecin 
were produced by high pressure homogenization.32 The in vitro release 
of the drug lasted for one week. After intravenous injections of 1.3 mg/kg camptothecin 
to mice, the drug residence time in the body was significantly prolonged by 
532 Kreuter 
the nanoparticles compared with the solution, and the plasma AUC was increased 
by a factor of 5, the brain AUC even by a factor of 10, and the AUC in other organs 
by a factor of between 2 (lungs) and 8.7 (heart). An increase of the dose to 3.3 mg/kg 
camptothecin (factor 2.5) in the nanoparticle formulation further increased the 
plasma AUC by 2.7, the brain AUC by 2.6, and the AUC in the other organs by 
2.9 times on the average. 
Incorporation of 3',5'-dioctanoyl-5-fluoro-2'-deoxyuridine into solid lipid 
nanoparticles also increased its brain uptake.33 After i.v. injection of the SLNs, its 
brain AUC was increased two-fold over the solution of this compound. 
Different types of solid lipid nanoparticles with a size of about 100 nm, consisting 
of emulsifying wax/Brij® 78 and out of Brij® 72/polysorbate 80, were made 
by Lockman et al34'35 and Koziara et al.36'37 and investigated in rat brain perfusion 
experiments. For both nanoparticle types, a statistically significant uptake 
was observed compared with [14C]-sucrose in rat brain perfusion experiments, 
suggesting central nervous system uptake of the nanoparticles.36 Perfusion of 
the nanoparticles did not induce any statistically significant changes in barrier 
integrity, membrane permeability, or facilitated choline transport.34 [3H]-thiamine 
was then bound to the emulsifying wax/Brij® 78 nanoparticles via a PEG-spacer 
(distearoylphosphatidyl-ethanolamine (DSPE)-PEG-NHS) to target the particles to 
the thiamine transporters in the brain.35 Although an association with the thiamine 
transporter occurred, no difference in the brain uptake was observed in BALB/c 
mice after i.v. injection between emulsifying wax/Brij® 78 nanoparticles with protruding 
PEG chains on the outside and nanoparticles with thiamine bound to the 
PEG chains. The emulsifying wax/Brij® 78 solid lipid nanoparticles were then 
loaded with paclitaxel and tested in the U-1118 and HCT-15 cell lines and by rat brain 
perfusion. Entrapment of paclitaxel in the solid lipid nanoparticles significantly 
increased its brain uptake and its toxicity towards the P-glycoprotein expressing 
tumor cells.37 
3.2. Influence of PEGylation on the biodistribution 
of nanoparticles 
Besides, by coating with surfactants, the body distribution may also be altered 
by covalent attachment of polyethylene glycol (PEG) chains to the nanoparticle 
surface (PEGylation). Like a number of surfactants such as poloxamine 908 and 
1508,23 the nanoparticle-surface-bound PEG chains can prevent the opsonization 
and rapid capture and removal of the nanoparticulate carriers by the cells of the 
reticuloendothelial system (RES), and consequently, can significantly prolong the 
blood circulation times of the particles.38-44 Calvo et al.41 showed in mice and rats 
that the 14C-concentration in different brain tissues was also significantly enhanced 
Nanoparticulate Carriers for Drug Delivery to the Brain 533 
after intravenous injection of PEGylated [14C]-poly[methoxy poly (ethylene glycol) 
cyanoacrylate-co-hexadecyl cyanoacrylate] nanoparticles ([14C]-PEG-PHDCA 
nanoparticles) in comparison to uncoated or poloxamer 908- or polysorbate 80- 
coated [14C]-poly (hexadecyl cyanoacrylate) nanoparticles ([14C]-PHDCA nanoparticles). 
Surprisingly, coating with polysorbate 80 and also with poloxamer 908 led 
to lower brain concentrations than uncoated particles in both species. In addition, 
a species-dependent influence of the surfactants on the brain concentrations was 
observed; in mice, the brain concentrations of the [14C]-PHDCA nanoparticles were 
higher after coating with polysorbate 80 than with poloxamer 908, whereas this 
order was reversed in rats. It is important to further note that after reduction of 
the nanoparticle dose, while maintaining the same total polysorbate concentration, 
higher [14C] brain levels were observed with the polysorbate 80-coated nanoparticles 
than with the PEGylated [14C]-PEG-PHDCA particles. The authors suggested 
that these higher brain concentrations were caused by a higher BBB permeability as 
a result of higher free blood polysorbate concentrations at the lower nanoparticle 
dose, and tried to support their assumption by another experiment injecting i.v. 
5% [14C]-sucrose in a 1% polysorbate solution in saline, which also led to higher 
[MC]-sucrose levels in the brain.41 However, this assumption that free polysorbate 
80 concentrations up to 1% leads to an increased BBB permeability resulting 
in a larger drug transport, is contradicted by pharmacological studies with 
drugs.18'19'21'22,25-27'45-49 In all of these studies, 1% polysorbate 80, containing drug 
solutions without nanoparticles that were used as controls, were unable to achieve 
any significant pharmacological effects. 
Interestingly, in the body distribution study of Calvo et al.il the brain concentration 
pattern was not mirrored in the other organs and tissues. The highest blood 
concentrations were obtained in mice and rats with the poloxamer 908-coated particles, 
followed by the PEG-PHDCA particles. The poloxamer 908-coated nanoparticles 
also yielded the lowest total uptake in the RES organs in rats but not in mice. 
In the latter, the PEG-PHDCA particles achieved the lowest total RES organ uptake. 
Similar results as those of Calvo et al.iX were also obtained with solid lipid 
nanoparticles, consisting of stearic acid (non-stealth SLN) or stearic acid, i.e. PEG 
2000 (stealth SLN), Epicuron® 200, and taurocholate sodium, after intravenous 
injection to rats43 and rabbits.44 Doxorubicin was bound to these particles using 
hexadecylphosphate as a counterion. In the rabbit study, the amount of the stealth 
agent stearic acid-PEG 2000 was systematically increased in 0.15% steps from 0% to 
0.45%.44 All nanoparticles achieved much higher and prolonged plasma concentrations 
than the doxorubicin solution. The increase in the stearic acid-PEG contents 
was mirrored by an increase and prolongation of the doxorubicin plasma concentrations. 
A comparable increase was observed in the brain reaching a doxorubicin 
concentration of 240ng/g after administration of 1 mg/kg doxorubicin. After only 
534 Kreuter 
6 hrs with the PEGylated solid lipid nanoparticles with the highest stearic acid- 
PEG content, doxorubicin was still detectable. No doxorubicin was found in the 
brain after administration of the doxorubicin solution. As in the abovementioned 
studies,27'28 the nanoparticles decreased the heart concentration, and in addition, 
the liver and other organ concentrations of the doxorubicin. 
The biodistribution of the PEGylated [14C]-PEG-PHDCA nanoparticles was 
also tested by Calvo et al.50 in DA/Rj rats with experimental allergic encephalitis 
(EAE) and compared with [14C]-PHDCA nanoparticles. The PEGylated nanoparticles 
achieved much higher brain and spinal cord concentrations than the normal 
particles. The concentration of the PEG-PHDCA nanoparticles was significantly 
higher in the pathological situation, where the BBB permeability was increased 
and was especially pronounced in the white matter. An enhanced macrophage 
infiltration with macrophages containing nanoparticles was observed at the EAE 
lesions, confirming earlier results of Merodio et al.51 after intraperitoneal injection 
of albumin nanoparticles. This transport within macrophages could augment the 
overall nanoparticle transport across the BBB. Coating of the non-PEGylated [14C]- 
PHDCA nanoparticles with poloxamine 908 resulted in very low and insignificant 
brain and spinal cord concentrations, although this surfactant again achieved very 
high nanoparticle plasma levels.50 Consequently, the PEGylated poly cyanoacrylate 
nanoparticles may represent promising brain drug delivery systems for neuroinflammatory 
diseases. 
4. Pharmacology 
As mentioned above, dalargin was the first drug that was transported across the 
BBB using the polysorbate 80-coated nanoparticles. Besides polysorbate 80, coating 
of the poly(butyl cyanoacrylate) nanoparticles with polysorbate 20,40, and 60 also 
enabled a transport of the nanoparticle-bound dalargin across the BBB, whereas 
coating with other surfactants such as poloxamers 184, 188, 338, 407, poloxamine 
908, Cremophor® EZ, Cremophor® RH 40, and polyoxyethylene-(23)-laurylether 
(Brij® 35) achieved no effectS24 (Table 1), clearly demonstrating the importance of 
the surface properties of the nanoparticles for brain drug delivery. 
Dalargin was then followed by other antinociceptive drugs such as the opioid 
loperamide45 and the Met-enkephalin kyotorphin,46 both showing similar effects. 
Unlike dalargin and kyotorphin, loperamide is not a peptide and is very lipophilic 
in contrast to these compounds. However, it is a strong Pgp substrate, and for 
this reason, it normally cannot cross the blood-brain barrier. In contrast to binding 
to poly(butyl cyanoacyrylate) nanoparticles, this drug was not able to induce 
any antinociceptive response after binding to polylactic acid nanoparticles, neither 
after coating with polysorbate 80 nor after preparation of the polylactic acid 
Nanoparticulate Carriers for Drug Delivery to the Brain 535 
Table 1 Maximal possible antinociceptive effect (MPE [%]) obtained 
after intravenous injection of dalargin-loaded surfactant-coated poly 
(butyl cyanoacrylate) nanoparticles and amount of apolipoprotein E 
(apo E) adsorbed on the surface of these particles in percent of the total 
amount of adsorbed plasma proteins. (Adapted from Kreuter et al.2i 
and from Luck70). 
Surfactant 
Uncoated 
Polysorbate 20 
Polysorbate 40 
Polysorbate 60 
Polysorbate 80 
Poloxamer 338 
Poloxamer 407 
Cremophor® EL 
Cremophor® RH40 
MPE [%] 
4.1 ± 1.0 
51 ±19 
61 ±41 
30 ±36 
89 ±22 
1.4 ±2.4 
8.1 ±2.9 
11.7 ±15.1 
23 ± 1 7 
Apo E adsorbed [%] 
0 
21.6 
29.7 
13.9 
14.6 
0 
0 
0 
0 
nanoparticles in the presence of this surfactant (unpublished results), although 
particles with a large variety of compositions with different release characteristics 
were manufactured.52,53 These observations clearly demonstrate that the ability of 
nanoparticles to enable a delivery of nanoparticles across the BBB, in addition to 
the surface properties, also depends on the core polymer. 
Tubocurarine normally also cannot cross the BBB. It induces epileptic spikes 
after direct intraventricular injection of tubocurarine into the brain. This drug was 
bound to the poly(butyl cyanoacrylate) nanoparticles47 and used in brain perfusion 
experiments in rats, in which the development of epileptic spikes in the EEC was 
recorded. Addition of the polysorbate 80-overcoated tubocurarine-loaded nanoparticles 
to the perfusate induced frequent severe spikes in the EEC that were comparable 
to direct intraventricular injection of the drug into the brain whereas a 
normal EEC was obtained after a solution of tubocurarine, the turbocurarine solution 
combined with polysorbate 80 or uncoated tubocurarine-loaded nanoparticles 
were added to the perfusate. 
The novel NMDAreceptor antagonists MRZ 2/576 (8-chloro-4-hydroxy-l-oxol,
2-dihydropyridazino[4,5-b]quinoline-5-oxide choline salt) is a potent but rather 
short-acting (5-15 min) anticonvulsant after intravenous administration.48 This 
short action is most likely caused by the rapid elimination of the drug from the 
central nervous system by efflux pump-mediated transport processes. Accordingly, 
these efflux processes can be inhibited by pretreatment with probenecid. Probenecid 
pretreatment prolongs the anticonvulsive action of MRZ 2/576 from about 15 min 
to 150 min. Intravenous injection of MRZ 2/576 bound to poly(butyl cyanoacrylate) 
nanoparticles coated with polysorbate 80, led to an even more prolonged duration 
536 Kreuter 
of the anticonvulsive activity in mice of up to 210 min, and in combination with 
probenecid up to 300 min.48,49 
In contrast to MRZ 2/576, the NMD A receptor antagonist MRZ 2/596 (8-chlorol,
4-dioxo-l,2,3,4-tetrahydropyridazino[4,5-b]quinoline choline salt) is not able to 
cross the BBB at all. However, again after binding to the polysorbate 80-coated 
nanoparticles, MRZ 2/596 also yielded similar anticonvulsive effects.49 
Two other drugs, amitriptyline,46 a tricyclic antidepressant, and valproic acid, a 
first line antiepileptic drug,54 were also bound to polysorbate 80-coated nanoparticles. 
While the brain AUC of amitriptyline was increased after intravenous injection 
of the polysorbate 80-coated nanoparticles which was accompanied by a reduction 
in serum AUC,46 no brain concentration increase was observable with valproic 
acid.54 
As mentioned above under Sec. 3.1, some indications exist that surfactantcoated 
nanoparticles or solid lipid nanoparticles may also increase the distribution 
of some drugs into the brain after oral administration,30,31,55 and may even lead 
to pharmacological effects.56 Coating of poly (butyl cyanoacrylate) nanoparticles 
with polysorbate 80 yielded antinociceptive effects with dalargin via the oral route, 
although these effects were not as pronounced but rather prolonged as after the i.v. 
injection. 
5. Brain Tumors 
Brain tumors, especially malignant gliomas, belong to the most aggressive human 
cancers. Despite numerous advances in neurosurgical operative techniques, adjuvant 
chemotherapy, and radiotherapy the prognosis for patients remains very 
unfavorable.57,58 These tumors are characterized by a rapid proliferation, diffuse 
growth, and invasion into distant brain areas, in addition to extensive cerebral 
edema and high levels of angiogenesis. Nevertheless, the disruption of the bloodbrain 
barrier (BBB) remains a local event, which is evident in the tumor core, but 
absent at its growing margins. For this reason, anticancer drugs can penetrate into 
necrotic tumor areas, while the drug levels in peritumoral regions were reported to 
remain low or non-detectable.59 
For this reason, very efficient anticancer drugs such as doxorubicin cannot cross 
the intact BBB and reach only the necrotic but not the peritumoral areas. As noted 
above (Sec. 3.1), however, this drug reached very high brain concentrations of about 
6 /xg/g, after binding to poly(butyl cyanoacrylate) nanoparticles. These nanoparticles 
were then tested in rats with intracranially implanted glioblastoma 101 / 8 . 5 8 In 
contrast to many experimental tumors such as RG 2 and 9L which are characterized 
by a nodular growth, this tumor has a stable monomorphous structure and 
shows the characteristic histological picture of aggressive glioblastomas with fast 
Nanoparticulate Carriers for Drug Delivery to the Brain 537 
diffuse growth in the brain parenchyma and a rather low tendency towards necrosis. 
Therefore, it is morphologically very similar to human glioblastomas. Doxorubicin 
bound to the polysorbate 80-coated poly(butyl cyanoacrylate) nanoparticles 
injected at a dose level of 1.5 mg/kg/day at days 2,5 and 8 after tumor transplantation, 
increased the mean survival time by 85% and repeatedly led to the survival 
of 20 to 40% of the animals for 180 days (from 8 repetitions of unpublished results). 
After this time, the animals were sacrificed, and the absence of tumors was demonstrated 
by histology in these animals. In contrast, the controls, empty polysorbate 
80-coated poly(butyl cyanoacrylate) nanoparticles, doxorubicin solution, doxorubicin 
solution plus polysorbate 80, and doxorubicin bound to poly(butyl cyanoacrylate) 
nanoparticles without polysorbate 80, led to no or much shorter increases in 
survival times or number of long-time survivers (Table 2). No indications of neurotoxicity 
were observable by histology with the nanoparticles. Also, the toxicity 
against other organs appeared to be reduced by binding to nanoparticles, in comparison 
to the doxorubicin solution. 
Brigger et al.i2 showed an accumulation of 14C-labelled PEG-PHDCA and 
PHDCA nanoparticles after intravenous injection to Fischer rats bearing an intracerebrally 
transplanted 9L glioblastoma. This accumulation was accompanied by 
a pronounced tumor retention effect. The tumor concentrations of the PEGylated 
nanoparticles were about 3 times higher than with the normal PHDCA particles, 
and about 5-6 times higher than in the adjacent brain areas. Interestingly, in the 
Table 2 Increases in survival times (1ST [%]) and long-term survivors (survival 100-180 
days) of rats with intracranially transplanted glioblastoma 101/8 after 3 intravenous injection 
of doxorubicin (1.5 mg/kg/day or 2.5 mg/kg/day) on days 2, 5, and 8 after tumor 
transplantation. (Adapted from Steiniger et alp). 
3 x 1.5mg/kg n 1ST [%] survival 
100-180 days 
Control 21 0 
Empty nanoparticles 13 0 n.s. 0 
Doxorubicin solution 23 54* 0 
Doxorubicin solution + polysorbate 80 22 65* 2 
Doxorubicin bound to nanoparticles 23 38* 2 
Doxorubicin bound to nanoparticles + polysorbate 80 23 85* 5 
3 x 2.5mg/kg 
Control 10 0 
Doxorubicin solution 8 88* 0 
Doxorubicin solution + polysorbate 80 8 108* 0 
Doxorubicin bound to nanoparticles 7 62* 0 
Doxorubicin bound to nanoparticles + polysorbate 80 9 169* 2 
*Statistically difference to controls (p < 0.05); n.s. not statistically different from control. 
538 Kreuter 
tumor-bearing rats, the brain concentrations in the areas adjacent to the tumor as 
well as in the controlateral brain hemisphere were also increased significantly compared 
with normal animals without tumor, indicating a generally higher permeability 
in the diseased animals. This was supported by co-injection of [3H]-sucrose 
together with the nanoparticles. The [3H]-sucrose level ratios between tumor, adjacent 
brain area, and the adjacent brain area obtained with the two types of nanoparticles 
were similar to the 14C-nanoparticle level ratios, and much lower levels again 
resulted without tumors.42 
Unfortunately, these nanoparticles did not increase the survival of Fisher rats 
bearing the same tumor, 9L, after intracranial transplantation.60 Biodistribution 
studies revealed that the binding of doxorubicin to the nanoparticles decreased the 
tumor accumulation of the particles by a factor of 2.5, which may be the cause for 
the lack of efficacy of these particles against 9L. 
The loss of wild type tumor suppressor genes like p53 function renders many 
tumors resistant to the induction of apoptosis by drugs such as doxorubicin.61 
Therefore, the delivery of wild type suppressor genes across the BBB is of enormous 
importance for the therapy with highly active chemotherapeutic drugs. The 
possibility of suppressor gene delivery into the brain with nanoparticles was evaluated 
in rats with an intracranially implanted F98 rat glioblastoma. Five days after 
tumor implantation, these rats received an intravenous injection of a 6-galactosidase 
reporter bound to poly(butyl cyanoacrylate) nanoparticles coated with polysorbate 
80. The animals were sacrificed 24, 48, and 72 hrs after nanoparticle injection, 
and a time dependent transport of the gene across the endothelial cells and glial cells 
was obtained, showing the strongest gene expression in the experimental tumors, 
whereas injection of naked control DNA did not render any expression at all.61 
6. Toxicology 
The acute toxicity of empty poly(butyl cyanoacrylate) nanoparticles as well as of the 
above poly(butyl cyanoacrylate) nanoparticle formulations, doxorubicin solution in 
saline, doxorubicin solution plus 1% polysorbate 80 in saline, doxorubicin bound 
to nanoparticles, and doxorubicin bound to nanoparticles coated with polysorbate 
80, was assessed by Gelperina et al.62 in normal and glioma 101 /8-bearing rats. 
Doses up to 400 mg/kg of empty nanoparticles did not cause any mortality within 
the period of observation (30 days), nor did they affect body weight or weight of 
internal organs after intravenous injection. Higher doses cannot be administered 
intravenously because of biological63 and technical limitations. No significant difference 
in toxicity occurred between the groups obtaining i.v. the four doxorubicin 
formulations in healthy as well as in the tumor-bearing animals. The results indicated 
that the toxicity of doxorubicin bound to nanoparticles is similar or may even 
be lower than that of free doxorubicin.62 
Nanoparticulate Carriers for Drug Delivery to the Brain 539 
In the above described chemotherapy study of Steiniger et a/.58 with doxorubicin 
bound to the polysorbate 80-coated poly(butyl cyanoacrylate) nanoparticles, 
similar toxicological results were obtained. Alimited dose-dependent systemic toxicity 
was found in the group treated with doxorubicin in saline. Autopsy of the whole 
body in healthy animals in this study revealed an empty gastrointestinal tract only 
in all animals treated with doxorubicin. The healthy animals treated with doxorubicin 
solution also showed slight signs of lung edema, which was confirmed 
by histology. These changes were not observed in animals treated with doxorubicin 
bound to the nanoparticles: indications of short-term neurotoxicity, such as 
increased apoptosis in areas distant from the tumor, increased expression of GFAP 
or ezrin on distant astrocytes or degenerative morphological changes of neurons, 
were entirely absent in treated animals on day 12, as well as in long-term survivors. 
In addition, there was no indication of chronic glial activation in areas distant from 
the tumor site in long-term surviving rats. Moreover, long-term survivors did not 
exhibit any obvious neurological symptoms.58 
7. Mechanism of the Delivery of Drug Across the 
Blood-Brain Barrier with Nanoparticles 
Presently, the mechanism of the delivery of drugs with nanoparticles across the 
BBB is not totally elucidated. A number of possibilities were suggested for this 
mechanism4'5'64: 
1. An increased retention of the nanoparticles in the brain blood capillaries combined 
with an adsorption to the capillary walls. This could create a higher concentration 
gradient that would enhance the transport across the endothelial cell 
layer, and as a result, the delivery to the brain. 
2. The polysorbate 80 used as the coating agent could inhibit the efflux system, 
especially P-glycoprotein (Pgp). 
3. A general toxic effect on the brain vasculariture, leading to the permeabilization 
of the brain blood vessel endothelial cells. 
4. A general surfactant effect characterized by a solubilization of the endothelial 
cell membrane lipids that would lead to membrane fluidization and an enhanced 
drug permeability through the blood-brain barrier. 
5. The nanoparticles could lead to an opening of the tight junctions between the 
endothelial cells. The drug could then permeate through the tight junctions in 
free, or together with the nanoparticles, in bound form. 
6. The nanoparticles may be endocytosed by the endothelial cells, followed by the 
release of the drugs within these cells and the delivery to the brain. 
540 Kreuter 
7. The nanoparticles with bound drugs could be transcytosed through the endothelial 
cell layer. 
All these mechanisms could also work in combinations .4'5,64 
Mechanisms 1 and 2 appear to be unlikely for the following reasons: if the 
drug-loaded nanoparticles would have merely created a high drug concentration 
gradient by adherence to the inner surface of the blood capillary walls (mechanism 
1), the diffusing drug would still have been subjected to the highly efficient 
efflux transporters in the membranes of the endothelial cells. Oh the other hand, if 
the polysorbate 80 would have inhibited these efflux transporters (mechanism 2), 
injection of polysorbate 80-coated empty nanoparticles 5 or 30 min before injection 
of dalargin, should also have induced antinociceptive effects, which was not 
observed in this case.65 The view that mechnisms 1 and 2 are unlikely are additionally 
supported by the brain perfusion experiments of Koziara et al.36 
Olivier et al.66 postulated that the enhanced drug transport across the BBB was 
caused by a toxic effect by the polysorbate 80-coated nanoparticles, resulting in the 
permeabilization or disruption of the blood-brain barrier (mechanisms 3 and/or 5). 
Pointing in the same direction, Calvo et al.il tried to explain the higher [14C]-sucrose 
levels that they observed in the brain after i.v. injection of 5% [14C]-sucrose in a 1% 
polysorbate 80 solution in saline with BBB permeabilization caused by unbound 
free polysorbate 80 present in the nanoparticle formulations (mechanism 4). However, 
both hypotheses can be refuted by the abovementioned experiment, where 
no antinociceptive effects were obtained after pre-injection of the polysorbate 
80-coated empty nanoparticles.65 In addition, no antinociceptive responses were 
observed after the injection of dalargin nanoparticles coated with other surfactants 
such as poloxamers 184, 188, 338, 407, poloxamine 908, Cremophor® EZ, 
Cremophor® RH 40, and polyoxyethylene-(23)-laurylether (Brij® 35),24 further outruling 
mechanism 4, a general membrane fluidization. This opinion that toxicity 
is not the mechanism for the nanoparticle-mediated drug transport across the BBB 
was also substantiated by the experiments of Sun et al.67 and of Koziara et al.36 Partial 
coverage of the particles by polysorbate 80 was sufficient for brain delivery,67 
and the brain perfusion experiments showed that the nanoparticles did not induce 
any statistically significant changes in barrier integrity, membrane permeability or 
facilitated choline transport.34 Finally, opening of the tight junctions as the underlying 
mechanism (mechanism 5) can be refuted by the findings that no major increase 
in the inulin spaces was observable in rat brain perfusion experiments.26 Additionally, 
electron microscopical studies also did not find any evidence for an opening 
of the tight junctions.65 
Therefore, the most likely mechanism appears to be mechanism 6, endocytotic 
uptake of the nanoparticles carrying the drug. This mechanism was already shown 
Nanoparticulate Carriers for Drug Delivery to the Brain 541 
in vitro in tissue cultures of brain endothelial cells of human, bovine, porcine, mice, 
and rat origin.26,68'69 At an incubation temperature of 37°C, a significant and rapid 
uptake was observed with the polysorbate 80-coated nanoparticles, whereas without 
coating, this uptake was minimal and it was inhibited at 4°C, a temperature at 
which phagocytosis does not occur, or after treatment with cytochalasin B, a potent 
phagocytic uptake inhibitor.69 
Mechanism 6 is further supported by the observation that, in contrast to the 
abovementioned surfactants, poloxamer 184 etc., besides polysorbate 80, polysorbates 
20,40, and 60 were also able to induce antinociceptive effects after the coating 
of dalargin-loaded nanoparticles and injection to mice.24 In addition, all 4 polysorbates, 
20,40,60 and 80, and not the other surfactants, were able to adsorb apolipoprotein 
E (apo E) on the surface of the nanoparticles after their incubation in blood 
plasma70 (Table 1). Kreuter et al.M then showed that the dalargin nanoparticles were 
also able to induce antinociceptive effects after the adsorption of apolipoproteins 
E and B. These effects were even much higher after polysorbate 80 pre-incubation. 
Therefore, the following scenario can be suggested: due to the polysorbate on their 
surface, the nanoparticles adsorb apolipoproteins E and/or B from the blood after 
injection. The particles, thus seem to mimic lipoprotein particles, and are taken 
up by the brain endothelial cells that express numerous lipoprotein receptors via 
receptor-mediated endocytosis. Since the efflux transporters are mainly located in 
the luminal membrane, the drug can then be transported into the brain by diffusion, 
after release from the very rapidly biodegrading20 nanoparticle polymer. It is also 
possible that the nanoparticles are transcytosed (mechanism 7), although no concrete 
evidence for this mechanism exists at present. The nanoparticles, therefore, 
seem to act as a "Trojan Horse". This hypothesis that drug transport via endocytotic 
uptake of the nanoparticles represents the underlying pathway was also supported 
by Sun et al.,67 Koziara et al.,36 and Gessner et al.71 Since the lipoprotein receptors are 
overexpressed in brain tumors,72 the above suggested scenario, lipoprotein receptor 
interaction, would also be an explanation for the good efficacy of the polysorbate 
80-coated doxorubicin-loaded poly(butyl cyanoacrylate) nanoparticles.58 
8. Summary 
A number of drugs that normally cannot cross the blood-brain barrier (BBB), or 
only in insufficient amounts, can be transported across this barrier after binding 
to polysorbate-coated poly(butyl cyanoacrylate) nanoparticles or to solid lipid 
nanoparticles, and achieve significant brain drug concentrations and pharmacological 
effects in the brain after intravenous injection. These drugs include the 
hexapeptide dalargin and the dipeptide kyotorphin, loperamide, tubocurarine, 
amitriptyline, the NMDA receptor antagonists MRZ 2/576 and MRZ 2/596, 
542 Kreuter 
doxorubicin, idarubicin, campthothecin, paclitaxel, as well as tobramycin. Doxorubicin 
bound to polysorbate 80-coated poly(butyl cyanoacrylate) nanoparticles was 
able to strongly improve the survival time of rats with intracranially transplanted 
glioblastoma 101 / 8 , an extremely aggressive tumor. The general toxicity of this drug 
was not increased by binding to the nanoparticles. PEGylation of poly cyanoacrylate 
nanoparticles prolonged their blood circulation time after intravenous injection 
and strongly increased their concentration in the intracranially transplanted 
glioblastoma 9L, but failed to prolong the survival of these rats. 
The mechanism of the nanoparticles-mediated drug transport across the BBB 
after intravenous injection seems to be the adsorption of apolipoproteins from the 
blood, leading to receptor-mediated endocytotic uptake of the particles into the 
brain capillary endothelial cells via lipoprotein receptors. The nanoparticles can 
then release the drugs within these cells, followed by the diffusion into the brain, 
or the access to the brain by transcytosis. 
9. Conclusions 
Poly cyanoacrylate nanoparticles or solid lipid nanoparticles (SLN) can enable the 
transport of many essential drugs across the blood-brain barrier (BBB) that normally 
cannot cross this barrier.4'5 The nanoparticles may be even useful for the delivery 
of larger and complex molecules such as proteins,18,19 nucleic acids and genes61 
across this barrier. They may also improve the treatment of brain tumors, since 
after binding to nanoparticles coated with polysorbates, anti-tumor drugs are also 
transported across the intact BBB,27,58 thereby accessing sites that cannot be reached 
by most anti-cancer drugs. 
Although the mechanism for the transport of nanoparticle-bound drugs across 
the BBB is not fully elucidated presently, binding of apolipoproteins after their 
injection into the blood stream, followed by receptor-mediated endocytotic uptake 
of the particles into the brain capillary endothelial cells, seems to be the most likely 
mechanism. Thus, the nanoparticles would act as a "Trojan Horse" which can then 
release the drugs within these cells, or after transcytosis into the brain. 
References 
1. Begley DJ (2004) Delivery of therapeutic agents to the central nervous system: The problems 
and possibilities. Pharmacol Ther 104:29^15. 
2. Brightman M (1992) Ultrastructure of the brain endothelium. Bradbury MWB (ed.) Physiology 
and Pharmacology of the Blood-Brain Barrier. Handbook of Experimental Pharmacology 
103 Springer: Berlin, Heidelberg, pp. 1-22. 
3. Begley DJ (1996) The blood-brain barrier: Principles for targeting peptides and drugs to 
the central nervous system. / Pharm Pharmacol 48:136-146. 
Nanoparticulate Carriers for Drug Delivery to the Brain 543 
4. Kreuter J (2001) Nanoparticulate systems for brain delivery of drugs. Adv Drug Del 
47:65-81. 
5. Kreuter J ((2002) Transport of drugs across the blood-brain barrier by nanoparticles. Curr 
Med Chem — Central Nervous Syst Agents 2:241-249. 
6. Begley DJ and Brightman MW (2003) Structural and functional aspects of the bloodbrain 
barrier. Prokai L and Prokai-Tatrai K (eds.) Peptide Transport and Delivery into the 
Central Nervous System. Progress in Drug Delivery. 61 Birkhauser Verlag: Basel, pp. 39-78. 
7. Rapoport SI (1996) Modulation of the blood-brain barrier permeability. / Drug Targ 3: 
417-425. 
8. Dehouck B, Fenart L, Dehouck M-P, Pierce A, Torpier G and Cecchelli R (1997) A new 
function for the LDL receptor: Transcytosis of LDL across the blood-brain barrier / Cell 
Biol 138:877-889. 
9. Gummerloch MK and Neuwelt EA (1992) Drug entry into the brain and its pharmacologic 
manipulation. Bradbury MWB (ed.) Physiology and Pharmacology of the Blood- 
Brain Barrier. Handbook of Experimental Pharmacology. 103 Springer: Berlin, Heidelberg, 
pp. 525-542. 
10. Remsen LG, Trail PA, Hellstrom I, Hellstrom KE and Neuwelt EA (2000) Enhanced 
delivery improves the efficacy of a tumor-specific doxorubicin immunoconjugate in a 
human brain tumor xenograft model. Neurosurgery 46:704-709. 
11. Pardridge WM, Buciak JL and Friden PM (1991) Selective transport of an anti-transferrin 
receptor antibody through the blood-brain barrier in vivo. J Pharmacol Exp Ther 259: 
66-70. 
12. Huwyler J, Wu D and Pardridge WP (1996) Brain drug delivery of small molecules using 
immunoliposomes. Proc Natl Acad Sci 93:14164-14169. 
13. Zhou X and Huang L (1992) Targeted delivery of DANN by liposomes and polymers. 
/ Control Rel 19:269-274. 
14. Chen D and Lee KH (1993) Biodistribution of calcitonin encapsulated in liposomes 
in mice with particular reference to the central nervous system. Biochem Biophys Acta 
1158:244-250. 
15. Kreuter J (1994) Nanoparticles. Swarbrick J and Boylan JC (eds.) Encyclopedia of Pharmaceutical 
Technology. 10 Marcel Dekker: New York, pp. 165. 
16. Birrenbach G and Speiser PP (1976) Polymerized micelles and their use as adjuvants in 
immunology. / Pharm Sci 65:1763-1766. 
17. H. Kopf H, Joshi RK, Soliva M and Speiser PP (1976) Studium der Mizellpolymerisation 
in Gegenwart niedermolekularer Arzneistoffe. 1. Herstellung und Isolierung der 
Nanopartikel, Restmonomerenbestimmung, physikalisch-chemische Daten. Pharm Ind 
38:281-284. 
18. Alyautdin R, Gothier D, Petrov V, Kharkevich D and Kreuter J (1995) Analgesic activity 
of the hexapeptide dalargin adsorbed on the surface of polysorbate 80-coated poly(butyl 
cyanoacrylate) nanoparticles. Eur } Pharm Biopharm 41:44^8. 
19. Kreuter }, Alyautdin RN, Kharkevich DA and Ivanov AA. (1995) Passage of peptides 
through the blood-brain barrier with colloidal polymer particles (nanoparticles). Brain 
Res 674:171-174. 
544 Kreuter 
20. Grislain L, Couvreur P, Lenaerts V, Roland M, Deprez-De Campenere D and Speiser P 
(1983) Pharmacokinetics and distribution of a biodegradable drug-carrier. Int } Pharm 
15:335-345. 
21. Schroeder U and Sabel BA (1996) Nanoparticles, a drug carrier system to pass the bloodbrain 
barrier, permit central analgesic effects of i.v. dalargin injections. Brain Res 710: 
121-124. 
22. Ramge P, Kreuter J and Lemmer B (1999) Circadian phase-dependent antinociceptive 
reaction in mice after i. v. injection of dalargin-loaded nanoparticles determined by the 
hot-plate test and the tail-flick test. Chronobiol Int 17:767-777. 
23. Troster SD, Miiller U and Kreuter 1(1990) Modification of the body distribution of 
poly(methyl methacrylate) nanoparticles by coating with surfactants. Int J Pharm 61: 
85-100. 
24. Kreuter J, Petrov VE, Kharkevich DA and Alyautdin RN (1997). Influence of the type 
of surfactant on the analgesic effects induced by the peptide dalargin after its delivery 
across the blood-brain barrier using surfactant-coated nanoparticles. / Control Rel 49: 
81-87. 
25. Schroeder U, Schroeder H and Sabel BA (2000). Body distribution of 3H-labelled 
dalargin bound to poly(butyl cyanoacrylate) nanoparticles after i.v. injections to mice. 
Life Sciences 66:495-502. 
26. Alyautdin RN, Reichel A, Lobenberg R, Ramge P, Kreuter J and Begley DJ (2001) Interaction 
of poly(butylcyanoacrylate) nanoparticles with the blood-brain-barrier in vivo and 
in vitro. } Drug Targ 9:209-221. 
27. Gulyaev AE, Gelperina SE, Skidan IN, Antropov AS, Kivman GY and Kreuter I (1999) 
Significant transport of doxorubicin into the brain with polysorbate 80-coated nanoparticles 
Pharm Res 16:1564-1569. 
28. Couvreur P, Kante B, Grislain L, Roland M and Speiser P (1982) Toxicity of polyalkylcyanoacrylate 
nanoparticles II: Doxorubicin-loaded nanoparticles. / Pharm Sci 71: 
790-792. 
29. Zara GP, Cavalli R, Fundaro A, Bargoni A, Caputo O and Gasco MR (1999) Pharmacokinetics 
of doxorubicin incorporated into solid lipid nanospheres (SLN). Pharmacol Res 
40:281-286. 
30. Zara GP, Cavalli R, Fundaro A, Bargoni A, Caputo O and Gasco MR (2002) Pharmacokinetics 
and tissue distribution of idarubicin-loaded solid lipid nanoparticles after 
duodenal administration to rats. ] Pharm Sci 91:1324-1333. 
31. Bargoni A, Cavalli R, Zara GP, Fundaro A, Caputo O and Gasco MR (2001) Transmucosal 
transport of tobramycin incorporated in solid lipid nanoparticles (SLN) after duodenal 
administration to rats. Part II - Tissue distribution. Pharmacol Res 43:497-502. 
32. Yang SC, Lu LF, Cai Y, Zhu JB, Liang BW and Yang CZ (1999) Body distribution in mice 
of intravenously injected camptothecin solid lipid nanoparticles and targeting effect on 
brain. / Control Rel 59:299-307. 
33. Wang JX and Zhang ZR (2002) Enhanced brain targeting by synthesis 3',5'-dioctanoyl-5- 
fluoro-2'-deoxyuridine and incorporation into solid lipid nanoparticles. Eur J Biopharm 
54:285-290. 
« 
Nanoparticulate Carriers for Drug Delivery to the Brain 545 
34. Lockman PR, Koziara J, Roder, KE, Paulson J, Abbruscato TJ, Mumper RJ and Allen DD 
(2003) In vitro and in vivo assessment of baseline blood-brain barrier parameters in the 
presence of novel nanoparticles. Pharm Res 20:705-713. 
35. Lockman PR, Oyewumi MO, Koziara J, Roder, KE, Paulson J, Mumper RJ and Allen DD 
(2003) Brain uptake of thiamine-coated nanoparticles. / Control Rel 93:271-282. 
36. Koziara MJ, Lockman PR, Allen DD and Mumper RJ (2003) In situ blood-brain barrier 
transport of nanoparticles. Pharm Res 20:1772-1778. 
37. Koziara MJ, Lockman PR, Allen DD and Mumper RJ (2004) Paclitaxel nanoparticles for 
the potential treatment of brain tumors. / Control Rel 99:259-269. 
38. Gref R, Minamitake Y, Peracchia MT, Trubetskoy V, Torchilin V and Langer R (1994) 
Biodegradable long-circulating polymeric particles Science 263:1600-1603. 
39. Bazile D, Prud'Homme C, Bassoullet M-T, Marlard M, Spenlehauer G and Veillard M 
(1995) / Pharm Sci 84:493^98. 
40. Peracchia MT, Fattal E, Deasmaele D, BesnardM, Noel JP, Gomis JM, Appel M, D'Angelo J 
and Couvreur P (1999) Stealth PEGylated polycyanoacrylate nanoparticles for intrave 
nous administration and splenic targeting. / Control Rel 60:121-128. 
41. Calvo P, Gouritin B, Chacun H, Desmaele D, D'Angelo J, Noel J-P, Georgin D, Fatal E, 
Andreux P and Couvreur P (2001) Long-circulating PEGylated polycyanoacrylate 
nanoparticles as new drug cariers for brain delivery. Pharm Res 18:1157-1166. 
42. Brigger I, Morizet J, Aubert G, Chacun H, Terrier-Lacombe M-J, Couvreur P and Vassal 
G (2002) Poly(ethylene glycol)-coated hexadecylcyanoacrylate nanospheres display a 
combined effect for brain tumor targeting. / Pharmacol Exp Ther 303:928-936. 
43. Fundaro A, Cavalli R, Bargoni A, Vighetto D, Zara GP and Gasco MR (2000) Non-stealth 
and stealth solid lipid nanoparticles (SLN) carrying doxorubicin: Pharmacokinetics and 
tissue distribution after i.v. administration to rats. Pharmacol Res 42:337-343. 
44. Zara GP, Cavalli R, Bargoni A, Fundaro A, Vighetto D and Gasco MR (2002) Intravenous 
administration of non-stealth and stealth doxorubicin-loaded solid lipid nanoparticles 
at increasing concentration of stealth agent: Pharmacokinetics and distribution in brain 
and other tissue. / Drug Targ 10:327-335. 
45. Alyautdin RN, Petrov VE, Langer K, Berthold A, Kharkevich DA and Kreuter J (1997) 
Delivery of loperamide across the blood-brain barrier with poly-sorbate 80-coated polybutylcyanoacrylate 
nanoparticles. Pharm Res 14:325-328. 
46. Schroeder U, Sommerfeld P, Ulrich S and Sabel BA. (1998) Nanoparticle technology for 
delivery of drugs across the blood-brain barrier. / Pharm Sci 87:1305-1307. 
47. Alyautdin RN, Tezikov EB, Ramge P, Kharkevich DA, Begley DJ and Kreuter J (1998) 
Significant entry of tubocurarine into the brain of rats by absorption to polysorbate 
80-coated polybutyl-cyanoacrylate nanoparticles: An in situ brain perfusion study. / 
Microencapsul 15:67-74. 
48. Friese A, Seiler E, Quack G, Lorenz B and Kreuter J (2000) Enhancement of the 
duration of the anticonvulsive activity of a novel NMDA receptor antagonist using 
poly(butylcyanoacrylate) nanoparticles as a parenteral controlled release delivery system. 
Eur } Pharm Biopharm 49:103-109. 
546 Kreuter 
49. Friese A. (2000) Kleinpartikulare Tragersysteme (Nanopartikel) als ein parenterales Arzneistofftransportsystem 
zur Verbesserung der Bioverfiigbarkeit ZNS-aktiver Substanzen dargestellt 
am Beispiel der NMDA-Rezeptor-Antagonisten MRZ 2/576 und MRZ 2/596. Ph.D. Thesis, 
JW Goethe-Universitat Frankfurt, Frankfurt. 
50. Calvo P, Gouritin B, Villarroya H, Eclancher F, Giannavola C Klein C, Andreux P 
and Couvreur P (2002) Quantification and localization of PEGylated polycyanoacylate 
nanoparticles in brain and spinal cord during experimental allergic encephalomyelitis 
in the rat. Eur J Neurosci 15:1317-1326. 
51. Merdio M, Irache JM, Eclancher F, Mirshadi M and Villarroya H (2000) Distribution 
of albumin nanoparticles in animals induced with the experimental allergic 
encephalomyelitis. / Drug Targ 8:289-303. 
52. Ueda M and Kreuter J (1997) Optimization of the preparation of loperamide-loaded 
poly(L-lactide) nanoparticles by high pressure emulsification-solvent evaporation. 
/ Microencapsul 5:593-605. 
53. Ueda M, Iwata A and Kreuter J (1998) Influence of the preparation methods on the drug 
release behaviour of loperamide-loaded nanoparticels. / Microencapsul 15:361-372. 
54. Darius J, Meyer FP, Sabel BA and Schroeder U (2000) Influence of nanoparticles on 
the brain-to-serum distribution and the metabolism of valproic acid in mice. / Pharm 
Pharmacol 52:1043-1047. 
55. Yang S, Zhu J, Lu Y and Yang C (1999) Body distribution of camptothecin solid lipid 
nanoparticles after oral administration. Pharm Res 16:751-757. 
56. Schroeder U, Sommerfeld P and Sabel BA (1998) Efficacy of oral dalargin-loaded 
nanoparticle delivery across the blood-brain barrier. Peptides 19:777-780. 
57. DeAngelis LM (2001) Brain Tumors. New Engl]Med 344:114-123. 
58. Steiniger SCJ, Kreuter J, Khalansky AS, Skidan IN, Bobruskin AI, Smirnova ZS, Severin 
SE, Uhl R, Kock M, Geiger KD and Gelperina SE (2004) Chemotherapy of glioblastoma 
in rats using doxorabicin-loaded nanoparticles. hit} Cancer 109:759-764. 
59. Donelli MG, Zucchetti M and D'lncalci M (1992) Do anticancer agents reach the tumor 
target in the human brain? Cancer Chemother Pharmacol 30:251-260. 
60. Brigger I, Morizet J, Laudani L, Aubert G, Appel M, Velasco V, Terrier-Lacombe M-J, 
Desmaele D, d'Angelo J, Couvreur P and Vassal G (2004) Negative preclinical results with 
stealthu nanosphere-encapsulated doxorubicin in an orthotropic murine brain tumor 
model. / Control Rel 100:29^0. 
61. Walz CM, Ringe K and Sabel BA. (2002) Nanoparticles in brain tumor therapy. Controlled 
Release Society 30th Annual Meeting Proc. Glasgow: # 630. 
62. Gelperina SE, Khalansky AS, Skidan IN, Smirnova ZS, Bobruskin AI, Severin SE, 
Turowski B, Zanella FE and Kreuter J (2002) Toxicological studies of doxorubicin bound 
to polysorbate 80-coated poly(butyl cyanoacrylate) nanoparticles in healthy rats and 
rats with intracranial glioblastoma. Toxicol Lett 126:131-141. 
63. Diehl K-H, Hull R, Morton D, Pfister R, Rabemampianina Y, Smith D, Vidal J-M and van 
de Vorstenbosch C (2001) A good practice guide to the administration of substances and 
removal of blood, including routes and volumes. / Appl Toxicol 21:15-23. 
Nanoparticulate Carriers for Drug Delivery to the Brain 547 
64. Kreuter J, Shamenkov D, Petrov V, Ramge P, Cychutek K, Koch-Brandt C and Alyautdin R 
(2002) Apolipoprotein-mediated transport of nanoparticle-bound drugs across the 
blood-brain barrier. / Drug Targ 10:317-325. 
65. Kreuter J, Ramge P, Petrov V, Hamm S, Gelperina SE, Engelhardt B, Alyautdin R, von 
Briesen H and Begley DJ (2003) Direct evidence that polysorbate 80-coated poly(butyl 
cyanoacrylate) nanoparticles deliver drugs to the CNS via specific mechanisms requiring 
prior binding of drugs to the nanoparticles. Pharm Res 20:409^116. 
66. Olivier J-C, Fenart L, Chauvet R, Pariat C, Cecchelli R and Couet W (1999) Indirect 
evidence that drug brain targeting using polysorbate 80-coated polybutylcyanoacrylate 
nanoparticles is related to toxicity. Pharm Res 16:1836-1842. 
67. San W, Xie C, Wand H and Hu Y (2004) Specific role of polysorbate 80 coating on the 
targeting of nanoparticles to the brain. Biomater 25:3065-3071. 
68. Borchard G, Audus KL, Shi F and Kreuter J (1994) Uptake of surfactant-coated 
poly(methyl methacrylate)-nanoparticles by bovine brain microvessel endothelial cell 
monolayers. Int} Pharm 110:29-35. 
69. Ramge P, Unger RE, Oltrogge JB, Zenker D, Begley D, Kreuter J and von Briesen H 
(2000) Polysorbate 80-coating enhances uptake of polybutylcyanoacrylate (PBCA)- 
nanoparticles by human, bovine and murine primary brain capillary endothelial cells. 
Eur J Neurosci 12:1931-1940. 
70. Luck M (1997) Plasmaproteinadsorption als moglicher Schltisselfaktor fur eine kontrollierte 
Arzneistoffapplikation mit partikularen Tragern. Ph.D. Thesis, Freie Universitat 
Berlin, pp. 14-24,137-154. 
71. Gessner A, Olbrich C, Schroder W, Kayser O and Miiller RH (2001) The role of plasma 
proteins in brain targeting: Species dependent protein adsorption patterns on brainspecific 
lipid drug conjugate (LDC) nanoparticles. Int} Pharm 214:87-91. 
72. Gutman RL, Peacock G and Lu DR (2000) Targeted drug delivery for brain cancer treatment. 
/ Control Rel 65:31-41. 
This page is intentionally left blank
25 
Nanoparticles for Targeting Lymphatics 
William Phillips 
1. Introduction 
Nanoparticles have received increasing attention as lymph node drug delivery 
agents.1-6 The desire to develop of new methods of lymph node drug delivery 
stems from the recent awareness of the importance of lymph nodes in cancer prognosis, 
their importance for vaccine immune stimulation and the realization that the 
lymph nodes harbor human immunodeficiency virus (HIV) as well as other infectious 
diseases. New methods of delivering drugs and antigens to lymph nodes are 
currently under investigation. 
The lymphatic system consists of a network of lymphatic vessels and lymph 
nodes that serve as a secondary vascular system to return fluid that has leaked 
from the blood vessels in the extremities and other organs back to the vasculature.7 
The lymphatic system also moves substantial volumes of fluid from the peritoneal 
cavity and pleural cavity back into the blood circulation. In addition to this critical 
role in the regulation of tissue fluid balance, the lymphatic system also plays an 
important role in intestinal absorption of fats and in the maintenance of an effective 
immune defense.7 Lymphatic vessels serve as a major transport route for the 
dissemination of antigens, microorganisms and tumor cells as well as interstitial 
molecules that have gained entry to the interstitial space.8 These lymphatic vessels 
are also traversed by immune cells such as dendritic cells, macrophages and as 
their name reveals, lymphocytes. As a part of the lymphatic system that recycles 
fluid from the interstitial spaces and the body's cavity back to the arteriovenous 549 
550 Phillips 
vascular system, the lymph nodes are ideally positioned to serve as surveillance 
organs to monitor microbial invasion and to defend the body against these invading 
microorganisms. The importance of the lymphatic system for the development of 
an effective immune response has led one author to describe the lymphatic vessels 
and lymph nodes as the body's "information superhighway".9 
1.1. The lymphatic vessels 
Lymphatic vessels are composed of thin, endothelial cell lined lymphatic capillaries 
located in spaces between cells and tissues. These lymphatic vessels are distributed 
throughout the body with the exception of cartilage, optic cornea and lens, and 
the central nervous system. Lymph fluid originating from the interstitial spaces 
between tissue cells and from within the body's cavities moves into lymphatic 
capillaries through lymph nodes and back into the blood circulation. The overlapping 
nature of the lymphatic endothelial cells and loose attachment of intercellular 
junctions allows for the absorption of interstitial fluid into the lymphatic capillaries. 
This mechanism also explains how macromolecules, infectious organisms and 
subcutaneously injected nanoparticles gain entrance into the lymphatic circulation. 
These lymphatic capillaries carry lymph fluid into collecting lymphatic vessels and 
channels which slowly flow in the afferent lymphatic vessels into the lymph node. 
After passing through the lymph node, the lymph fluid then exits the lymph node 
through the efferent lymphatic vessels. The lymph fluid flows slower and under 
much lower pressure than blood in the artery and veins. Its flow rate can be greatly 
accelerated by body movement. The efferent vessels combine to form lymphatic 
vessels that branch either to the next set of lymph nodes or to larger lymphatic 
trunks as illustrated in Fig. 1. In this way, lymph fluid of different organs and the 
body's extremities in addition to body cavities is collected by large lymphatic trunks 
which feed into one of the two lymphatic ducts: the thoracic duct and right lymphatic 
duct. From these ducts, the lymph fluid then returns to the blood stream 
through veins in the neck region (i.e. internal jugular and subclavian veins).1011 
The lymphatic system also returns fluid from the body's cavities back to the 
blood stream. This includes fluid from the pleural space surrounding the lungs, 
the peritoneal space surrounding the intestines, the articular cavity of the joints, 
and the central spinal fluid surrounding the brain. The rate of movement of fluid 
from these body cavities varies between different cavities. However, the volume 
of fluid moving through the lymphatic vessels coming from the body's cavities is 
substantial. Studies in conscious sheep with tracers have found that clearance of 
fluid from the peritoneal cavity averaged 2.4ml/hr per kg of body weight, which 
was more than twice as high as when the sheep were anesthetized.12 This rate of 
Nanoparticles for Targeting Lymphatics 551 
Mediastinal 
Lymph N 
Thoracic 
Duct 
Pleural 
Cavity 
Peritoneal 
Cavity 
Diaphragm 
Fig. 1. The lymphatic system includes lymphatic vessels draining from the extremities and 
head and neck region as well as the fluid moving from the cavities of the body. It is estimated 
that 400-600 lymph nodes that filter drainage from the lymph vessels are in the average 
human body. 
fluid movement in a 70-kg human would be more than 165 ml per hr or nearly 
4 liters per day. 
1.2. Lymph nodes 
The lymphatic system is also interspersed with lymph nodes placed at intervals 
along this lymphatic vessel network as shown in Fig. 1. Lymph nodes are encapsulated 
dense masses of lymphoreticular tissue situated along the pathway of 
drainage of the lymph. There are estimated to be 400-600 lymph nodes in the human 
552 Phillips 
Fig. 2. This diagram illustrate the structure of the lymph node. Efferent lymphatic vessels 
delivery lymph fluid to the lymph node and afferent lymphatic vessels take the lymphatic 
fluid from the lymph node. Each lymph node is supplied by an artery and a vein. Lymphatic 
fluid is filtered through the sinuses of the lymph nodes that are lined with macrophages to 
phagocytized foreign particulate agents. Lymph nodes also contain cortical, paracortical and 
medullary regions which contain different immune system cells. 
body. The outer capsule of the lymph node is composed of dense collagenous fibers 
and smooth muscle fibers. The interior of the lymph nodes is organized into different 
functional zones populated by different sorts of lymphocytes, as well as 
accessory and stromal cells.11,13 These lymph node zones as shown in Fig. 2 are: 
• The cortex, which includes the lymphoid follicles with their germinal centers. 
This is the B-cell area of the lymph node which is associated with humoral 
immune mechanisms. 
• The paracortex, which is the densely cellular area that extends between the lymphoid 
follicles. This is the T-cell area which is the main site of cellular immunity. 
• The sinuses, a complex system of channels where macrophages belonging to the 
mononuclear phagocytic system (MPS) reside. 
• The medulla, rich in sinuses where the main site of plasma cell proliferation and 
production of antibodies (the medullary cords) are located. 
The three main functions of the lymph nodes are the formation of lymphocytes 
known as lymphopoiesis, lymph filtration, and antigen processing.9,14 In terms of 
Nanoparticles for Targeting Lymphatics 553 
lymph fluid filtration, the lymph nodes provide two main types of filtration: a simple 
mechanical type through the reticular meshwork which traverses the sinuses; 
and a phagocytic filtration by macrophages and reticular cells, which is aided by 
the slow passage of the lymph fluid through the channels of the sinuses. 
One of the major functions of the lymph nodes is to help defend the body against 
diseases by filtering bacteria and viruses from the lymph fluid, and to support the 
activities of the lymphocytes, which furnish resistance to specific disease-causing 
agents. However, in abnormal conditions, as in the case of cancer and some infections, 
it is well known that lymph nodes can act as holding reservoirs from where 
tumor cells, bacteria or viruses can spread to other organs and regions of the body.7,11 
For example, in the case of cancer, disseminating tumor cells can take root in lymph 
nodes and form residual metastatic tumors that are difficult to detect and treat. 
In anthrax infection, endospores from Bacillus anthracis that gain entrance into 
the body are phagocytozed by macrophages and carried to regional lymph nodes, 
where the endospores germinate within the macrophages and become vegetative 
bacteria. The vegetative bacteria are then released from the lymph nodes, multiply 
in the lymphatic system and invade the blood stream causing massive septicemia.15 
Adequate therapy of lymph nodes affected by disease runs a considerable risk 
of side effects. For example, current methods of treating or preventing metastasis in 
lymph nodes are characterized by serious drawbacks including: (1) radical surgical 
excision of lymph nodes is a burdensome procedure, and the risk of postoperative 
lymph node cancer recurrence is often high; (2) external radiation therapy can 
damage sensitive organs unnecessarily, while delivering a small percentage of radiation 
to the targeted lymph nodes; and (3) intravenous chemotherapy to patients 
with advanced disease is associated with significant toxicity, even though adequate 
therapeutic concentrations in the targeted lymph nodes are rarely achieved.1 For 
these reasons, nanoparticle targeted drug carriers offer a potential solution to the 
challenge of adequate lymph node therapy. 
2. Potential for Nanoparticles for Drug Delivery 
to Lymphatics 
Nanoparticles are ideal structures for delivering therapeutic agents to the lymph 
nodes. Their ideal features are based on their size, which prevents their direct 
absorption into the blood, the large amount of drugs and other therapeutic agents 
that nanoparticles can carry, and their ability to be retained in the lymph nodes. In 
comparison, small molecules will be directly absorbed into the blood at the site of 
injection and will not move into the lymphatics. Although larger molecules such as 
dextran and albumin will move into the lymphatics, they rapidly pass through the 
draining lymph nodes and are not well retained in individual lymph nodes.16'17 
554 Phillips 
Although nanoparticles are too large to be directly absorbed into the blood 
stream, they are small enough to enter the lymph vessels and lymph nodes, following 
either subcutaneous injection, intradermal injection, intramuscular injection or 
injection directly into organs or tumors and injection into the body's cavities. Following 
subcutaneous injection or injection directly into the tissue of a body organ, it 
appears that a certain portion of nanoparticles are taken up locally and retained for 
a prolonged time, while another portion of the nanoparticles are cleared from this 
local site and move into the lymphatic vessels, where they can be trapped in lymph 
nodes or else move completely through the lymphatic system and return to the 
blood at the thoracic duct. Nanoparticles that are injected directly into the body's 
cavities appear to have much less local retention at the site of injection, as they 
disperse freely throughout the whole cavity and then drain almost completely into 
lymphatic vessels, where they can be trapped in lymph nodes or return to the blood 
circulation. These intracavitary sites whose fluid is cleared through the lymphatics 
include the pleural space surrounding the lungs, the peritoneal space surrounding 
the intestines, the articular cavity of the joints, and the central spinal fluid surrounding 
the brain. If the particular particle that is injected into the intracavitary space 
is only minimally retained in all of the draining lymph nodes, then the ultimate 
fate of the nanoparticles injected into a cavity can appear very similar to the same 
distribution that it would have had, following intravenous administration.18 This 
similar distribution has been demonstrated with radiolabeled nanoparticles that are 
injected into the peritoneal space. By 24 hrs following intraperitoneal administration, 
these nanoparticles have a high liver and spleen uptake and minimal retention 
in the peritoneum, as if they had been injected intravenously.18 This is the normal 
distribution of neutral and anionic liposomes unless there is special modification 
of the liposomes to increase their uptake in the lymph nodes.3,18 
Considering the importance of the lymphatics in relationship to many disease 
processes, the number of studies investigating drug delivery or targeting of other 
therapeutic agents to the lymphatics has been relatively modest.19 The recent development 
of an increasing number of different types of nanoparticles that can facilitate 
the lymphatic transport of therapeutic agents provides many new approaches to 
lymphatic drug delivery and the basic investigation of lymphatic transport. 
3. Importance of Lymph Nodes for Disease Spread and 
Potential Applications of Lymph Node Drug Delivery 
3.1. Cancer 
The majority of solid cancers spread primarily by lymph node dissemination.20 The 
status of the lymph node in regard to cancer metastasis is a major determinant of the 
Nanoparticles for Targeting Lymphatics 555 
patient's prognosis. This includes very small metastasis detected by histopathologic 
analysis of removed lymph nodes, as well as by magnetic resonance imaging (MRI) 
imaging using magnetic nanoparticle MRI contrast agents.21 Accurate lymph node 
staging is the most important factor that determines the appropriate care of the 
patient.22 Therapeutic interventions that treat metastatic cancer in lymph nodes 
with either surgery or local radiation therapy have been shown to improve patient 
survival.23 
3.2. HIV 
Primary infection with human immunodeficiency virus (HIV) is characterized by 
an early viremia, followed by a specific HIV immune response and a dramatic 
decline of virus in the plasma.24 Long after the HIV virus can be found in the 
blood, HIV can be found in high levels in mononuclear cells located in lymph 
nodes. Viral replication in these lymph nodes has been reported to be 10-100-fold 
higher than in the peripheral blood mononuclear cells.25 Drug delivery to these 
lymph node mononuclear cells is difficult with standard oral or intravenous drug 
administration. Although highly active antiretroviral therapy (HAART) reduces 
plasma viral loads in HIV infected patients by 90%, active virus can still be isolated 
from lymph nodes even after 30 months of HAART therapy. 
3.3. Filaria 
Lymph nodes are an important part of the life cycle of several parasite organisms, 
including filaria. Adult worms are found in the lymphatic vessels and lymph nodes 
of infected patients. These adult filaria are responsible for the obstruction of lymphatic 
drainage that causes swelling of extremities that are distal to the infected 
lymph node. The condition associated with very swollen limbs often found in 
patients with filarial disease has been termed elephantiasis. Ultrasound imaging 
can be used to visualize the adult worms by detecting the classic "filaria dance 
sign" which is associated with the adult worms.26 This ultrasound imaging has 
demonstrated that the worms reside in "nests" located in the lymph nodes and 
lymphatic vessels. The preferred site for adults worms in males is the intrascrotal 
juxtatesticular lymphatic vessels.27 Adult worms are also commonly found in the 
inguinal nodes28 and they have even been reported to be found in the mediastinal 
lymph nodes.29 
It is very difficult to eradicate the adult worms located in the lymphatic system, 
although microfilaria that are released into the blood stream from adult worms are 
very responsive to anti-filarial medications. This difficulty in treatment may be due 
to the localization of adult worms within "nests" in the lymphatic system, where 
556 Phillips 
drug penetration is very poor. Frequently, eradication of adult worms is not possible 
and it commonly takes a very extended course of medical therapy to have 
any effect on the adult worms.30 Nanoparticle drug delivery has potential for drug 
delivery in filarial disease, particularly before the lymphatics have become totally 
obstructed. Many asymptomatic patients have been shown to carry adult worms in 
their lymph nodes. Thus, early diagnosis may be crucial for the treatment of filaria 
before adult worms are well established in the lymphatic vessels and nodes. 
3.4. Anthrax 
New methods of treating anthrax have become of urgent interest, following the 
recent outbreak of terrorist caused infections and deaths in the United States as a 
result of terrorism. Following inhalation of the anthrax spores and their deposition 
in the lungs, the bacteria spread to the mediastinal lymph nodes where their 
local invasion and associated toxin production is the cause of death. Patients are frequently 
found to have a widened mediastinum due to the expansion of mediastinal 
lymph nodes with anthrax.31 
Computed tomography of the chest has been performed on 8 recent patients 
infected with inhalational anthrax. Mediastinal lymphadenopathy was present in 
7 of the 8 patients.32 In a recent case report of one patient, the anthrax bacillus was 
shown to be rapidly sterilized within the blood stream after initiation of antibiotic 
therapy. However, viable anthrax was still present in postmortem mediastinal 
lymph node specimens.33 This case demonstrates the difficulty that drugs have in 
penetrating the mediastinal lymph nodes. A potential use of nanoparticles could 
be for delivery of anti-anthrax drugs to the mediastinal lymph nodes for therapy 
or prevention of anthrax extension to the lymph nodes. 
3.5. Tuberculosis 
The tuberculosis infection is caused by mycobacteria which invade and grow chiefly 
in phagocytic cells. Tuberculosis is frequently found to spread from the lungs 
to lymph nodes, so that lymph node tuberculosis is the most comm on form of 
extrapulmonary tuberculosis. In one study, 71% of the tuberculosis lymph node 
involvement was located in the intrathoracic lymph nodes, while 26% of the cervical 
lymph nodes were involved with tuberculosis and 3% in the axillary lymph 
nodes.34 The development of methods to target drugs to these lymph nodes could 
greatly improve the therapy of tuberculosis and potentially decrease the amount of 
time that drug therapy is required. Currently, patients with tuberculosis are required 
to take medication for > 6 months. One possible reason for this lengthy treatment 
is the difficulty in deliverying drugs into these tubercular lesions. This requirement 
of lengthy drug treatment could also be responsible for the development of 
Nanoparticles for Targeting Lymphatics 557 
resistance to anti-tuberculosis drugs, as the organisms are exposed to relatively 
low levels of drugs over a very prolonged time. The development of resistance to 
anti-tuberculosis therapy is a growing health problem and the number of tuberculosis 
cases has been increasing worldwide. Multidrug-resistant tuberculosis presents 
an increasing threat to global tuberculosis control.35 Nanoparticles could be used 
to specifically carry high levels of drugs to lymph nodes containing tuberculosis. 
Nanoparticles encapsulating anti-tuberculosis drugs have already been developed 
as potential intravenous therapeutic agents for the treatment of tuberculosis.36 
3.6. Importance of lymph node antigen delivery for 
development of an immune response 
The importance of the lymph nodes in the development of an immune reaction 
induced by vaccines is gradually becoming recognized. Experimental evidence suggests 
that the induction of immune reactivity depends on antigen reaching and 
being available in lymphoid organs in a dose- and time-dependent manner.37 This 
concept has been termed the geographical concept of immune reactivity.37-40 The 
delivery of antigen to a lymph node in a manner that resembles an actual microbial 
invasion may be one of the most important functions of a vaccine adjuvant. The 
adjuvants are considered effective, if they either enhance or prolong expression of 
antigen components to reactive T cells in lymph nodes.38 Antigen-presenting cells 
are thought to be of critical importance in transporting antigen from the periphery 
to local organized lymphoid tissue. However, delivery of antigen to the lymph 
node by any means may be more important. Several studies have investigated the 
immune response following direct injection of antigen into lymph nodes. Instead 
of injecting peptide-based vaccines subcutaneously or intradermally, researchers 
injected these agents directly into the lymph nodes.39 This intralymphatic injection 
enhanced immunogenicity by as much as 106 times when compared with subcutaneous 
and intradermal vaccination. Intralymphatic administration induced CD8 
T cell responses with strong cytotoxic activity and interferon (IFN)-gamma production 
that conferred long-term protection against viral infections and tumors. This 
greatly increased response based on direct delivery to the lymph node has also been 
reported with naked DNA vaccines. Naked DNA vaccines are usually administered 
either intramuscularly or intradermally. When naked DNA was injected directly into 
a peripheral lymph node, immunogenicity was enhanced by 100- to 1000-fold, inducing 
strong and biologically relevant CD8(+) cytotoxic T lymphocyte responses.41 
Nanoparticles can be used to greatly increase the delivery of an antigen to the 
lymph node.39 For instance, animal experiments have shown that immunization 
by the intramuscular or the subcutaneous route with liposome-entrapped plasmid 
DNA encoding the hepatitis B surface antigen leads to much greater humoral (IgG 
558 Phillips 
subclasses) and cell mediated (splenic IFN-gamma) immune responses than with 
naked DNA.40 In other experiments with a liposome encapsulated plasmid DNA 
encoding a model antigen (ovalbumin), a cytotoxic T lymphocyte (CTL) response 
was also observed. These results could be explained by the ability of liposomes to 
protect their DNA content from local nucleases and direct it to antigen presenting 
cells (APCs) in the lymph nodes draining the injected site. 
In spite of this awareness of the importance of antigen delivery to lymph nodes, 
there have been very few studies in which the biodistribution of an injected vaccine 
antigen has been determined, following either subcutaneous, intradermal or intramuscular 
administration. Studies of antigen encapsulated within nanoparticles 
could easily be carried out using scintigraphic tracers and imaging. Scintigraphic 
imaging can provide quantitative information of the total dose and percentage of 
the antigen that reaches the lymph node. It appears that the purpose of a vaccine 
adjuvant is to simulate as closely as possible the delivery of a virus, that enters 
the body subcutaneously or through a body cavity, to the lymph node. One reason 
that nanoparticles appear to be useful for vaccine delivery is because their processing 
resembles that of viruses. Viruses can be considered as naturally occurring 
nanoparticles, against which the human immune system has evolved a defense 
mechanism. 
In this regard, it is remarkable that there have been very few studies investigating 
the distribution of nanosized viruses, following their administration subcutaneously 
or intracavitary in methods in which the first encounter with the immune 
system is likely to be in the lymph node. These studies could easily be performed 
in experimental small animal imaging models, by labeling the viruses with a scintigraphic 
imaging agent and injecting them subcutaneously or into a body cavity. 
Their distribution in the body could be followed by performing serial imaging studies. 
There have been, however, several studies of viral sized radiolabeled colloidal 
particles being injected subcutaneously, in which assumptions were made about 
the likely distribution of viral particles. In one of these studies, subcutaneously 
injected, viral-sized particles were found to initially arrive in the blood and later 
in the lymph.42 Accumulation in lymph and blood increases for a prolonged time 
following subcutaneous administration. The results of this study suggested the possibility 
that strategies could be developed to limit the spread of infectious agents 
by early aggressive local antiviral treatment. 
Nanoparticles also have the potential to be delivered to lymph nodes by means 
other than subcutaneous injection. Particles as large as 1.1 um in diameter have 
been found to be translocated from the nasal mucosa to lymph nodes following 
intranasal administration. 24 hrs after intranasal administration of relatively large 
1.1 um diameter fluorescent microspheres, significant fluorescence was visualized 
in the posterior cervical lymph nodes and in the mediastinal lymph nodes.43 
Nanoparticles for Targeting Lymphatics 559 
4. Factors Influencing Nanoparticle Delivery to Lymph Nodes 
4.1. Nanoparticle size 
Many factors appear to influence the fraction of the nanoparticles that are retained 
at the initial site of subcutaneous injection. Nanoparticle size appears to be one of 
the most important factors affecting the clearance of nanoparticles from the subcutaneous 
site of injection.5 The larger the size of the nanoparticles that are injected 
subcutaneously, the greater the fraction of the nanoparticles that will be retained 
locally and the lesser that will enter the lymphatic vessels and have a chance to 
target the lymph nodes.5'44 
Much work has been performed evaluating the effect of particle size of liposomal 
subcutaneously injected nanoparticles on lymph node targeting. Liposomes 
are nanoparticles composed of naturally occurring phospholipids that form spontaneously 
in an aqueous environment. Much recent research has investigated the 
potential of liposomes as carriers of drugs and other therapeutic agents to lymph 
nodes. When small neutral liposomes are injected subcutaneously, more than 60- 
70% of the liposomes will be cleared from the injection site by 24hrs,16,45 with only 
30-40% of the injected dose remaining at the site of injection. Liposomes larger 
than 500 nm will have 60-80% remaining at the injection site.16-45 The dose of lipid 
administered does not appear to have an effect on the percentage of liposomes 
retained in the lymph node. Lymph node uptake did not appear to become saturated 
over a large of lipid dose administered ranging from 10 nmol lipid to 10,000 
nmol of lipid.45 
Factors that enhance clearance of liposomes from a local site of subcutaneous 
injection also appear to decrease liposome uptake in the lymph node. For instance, 
larger liposomes are not cleared from the subcutaneous site of injection as readily 
as smaller liposomes; however, they are better retained in the lymph node. Even 
though larger liposomes are less well cleared from the injection site, their total 
retention in the lymph node is similar to other liposomes due to their improved 
lymph node retention. This improved lymph node retention by liposomes that are 
poorly cleared from the injection site results in liposome retention doses that are 
approximately equal to liposomes that have improved clearance from the local 
subcutaneous injection site.16 
4.2. Nanoparticle surf a ce 
Several studies have been carried out to determine the ideal nanoparticle surface 
characteristics for the delivery of drugs to the draining lymph nodes following subcutaneous 
injection. Moghimi et al. have performed studies of 45 nm polystyrene 
nanospheres which have been coated with poloxamers.4,46 The effect of a variety of 
560 Phillips 
different polyoxamers with varying lengths of ethylene oxide (EO) units has been 
studied. If the nanospheres are coated with polyoxamers with a large number of 
EO units per chain, the nanospheres will clear rapidly from the site of injection, but 
they will also escape removal by macrophages in the lymph nodes, so that lymph 
node uptake will be minimal. If nanospheres are not coated with any polyoxamer 
surface, they will remain largely retained at their site of subcutaneous injection 
and only a small percentage (< 3% of the injected dose) will accumulate in the 
lymph node. The polystyrene nanospheres with the most effective delivery and 
retention in the draining lymph nodes have a polyoxamer coating of 4-15 EO units 
per chain and a coating thickness of less than 3 nm.4 Nanospheres of this type have 
both rapid clearance from the subcutaneous site of injection, as well as significant 
retention in the draining lymph nodes. By 6 hrs, these ideal nanospheres have less 
than 50% located in the rat footpad and approximately 20% retained in the primary 
lymph node and 14% retained in the secondary node.4 Based on findings with the 
model particles, Moghimi suggested that it should be possible to develop liposomes 
or other nanoparticles with ideal properties for lymph node delivery. A suggestion 
was made to test liposomes composed of 5-7mol% of 2000 MW polyethylene 
glycol lipids.4 
Surface modification of liposomes with polyethylene glycol (PEG) did not 
appear to have a very large effect of lymph node uptake. Ousseren et al. found that 
the amount of liposomes that cleared from the injection site was slightly greater 
with the PEG-coated liposomes47; however, this improved clearance did not result 
in improved lymph node retention, because the fraction of PEG-liposomes retained 
by the lymph node is decreased. The slightly improved clearance of PEG-coated 
liposomes from the subcutaneous site of injection was also found by our research 
group.16 
4.3. Effect of massage on lymphatic clearance of subcutaneously 
injected liposomes 
The rate of clearance of nanoparticles from a subcutaneous injection site can be 
greatly accelerated with local manual massage.48 Without any mechanical stimulation, 
subcutaneously injected 200 mm liposomes are usually trapped in the interstitial 
subcutaneous space for a prolonged time. However, 5 min of manual massage 
over the subcutaneous injection site can clear up to 40% of the injected liposomes 
from the subcutaneous site into the blood via the lymphatic pathway. Investigators 
were able to use this effect to control the rate of drug delivery of the vasoconstricting 
hormone angiotensin II encapsulated in a liposome. They demonstrated that a 
physiological response to encapsulated drug (average blood pressure increase) can 
also be induced and modulated by massage.48 
Nanoparticles for Targeting Lymphatics 5 61 
4.4. Macrophage phagocytosis 
It is generally accepted that nanoparticles are retained in the lymph node by 
macrophage phagocytosis. Several research findings using nanoparticle liposomes 
appear to support this contention. For example, inclusion of phosphatidylserine 
(PS) in the liposome lipid formulation moderately increased lymph node uptake.45 
PS is a strong signal for stimulating macrophage uptake because it is present on 
the outer surface of cells undergoing apoptosis' instead of its usual location on the 
inner surface of the cell membrane.49 
The strong supporting evidence of the role of macrophages in lymph node 
uptake was provided by a study in which macrophages were temporarily depleted 
from lymph nodes, by prior administration of liposomes containing dichloromethylene 
diphosphonate (clodronate). Clodronate is toxic to macrophages and much 
previous work has been performed using clodronate to temporarily deplete 
macrophages in the liver.50,51 Six days after injection of the clodronate liposomes, 
small and large sized liposomes were also injected subcutaneously. There 
was a drastic reduction in the uptake of both large and small liposomes in the 
lymph node.52 This reduction in liposome uptake supports the hypothesis that 
macrophages play the most important role in nanoparticle uptake in lymph nodes. 
4.5. Fate of nanoparticles in lymph nodes 
Only a few studies have looked at the fate of nanoparticles once they arrive at the 
lymph node.5'53 In one study, subcutaneously injected liposomes were found to 
have accumulated in the subcapsular sinus. Subsequently, these liposomes were 
dispersed throughout the lymph node either by permeation along the sinus or 
within cells involved in liposome uptake such as macrophages. Once they were in 
the macrophages, the liposomes were observed to be digested by lysosomes.53 
5. Nanoparticle Diagnostic Imaging Agents for Determining 
Cancer Status of Lymph Nodes 
5.1. Subcutaneous injection of iodinated nanoparticles for 
computed tomography imaging 
Nanoparticles have been developed for the delivery of image contrast agents to 
lymph nodes. Lessons learned in the development of these nanoparticles as lymph 
node contrast agents can be applied to lymph node drug delivery. Subcutaneously 
injected iodinated nanoparticles were found to target the lymphatics and have been 
investigated as computed tomography imaging (CT) contrast agents.54-57 These 
nanoparticles were composed of ethyl ester of diatrizoioc acid, stabilized with 
562 Phillips 
3.5% Tetronic 908 with an average particle size of 250 nm. Nanoparticle accumulation 
in the draining lymph node was found to consistently enhance the contrast by 
at least 100 Hounsfield units in rabbits. A Hounsfield unit is a description of relative 
attenuation of X-rays in CT imaging. In this system, water density attenuation is 
assigned 0 and air density attenuation is —1000 and compact bone attenuation is 
1000. A contrast change of 100 Hounsfield units is easily recognized visually on the 
image. The contrast enhancement by the iodinated nanoparticles provided excellent 
detailed images of the intranodal lymph node architecture which would permit 
the diagnosis of cancer metastasis. 
Different size iodinated nanoparticles have been compared for lymph node 
targeting characteristics. When a comparison was made between the lymph node 
contrast of smaller 116 nm iodinated nanoparticles and larger 250 ran particles, there 
was minimal difference in the total lymph node contrast eventually obtained with 
the iodinated nanoparticles. However, the smaller nanoparticles were found to have 
somewhat faster kinetics for lymph node accumulation and they also clear more 
rapidly from the lymph node.58 
This iodinated contrast agent has been shown to aid in the discrimination of 
lymph nodes with cancer. In this study, perilesional subcutaneous injections (2 ml 
per lesion) of a 15% wt/vol iodinated nanoparticle suspension were made in pigs 
with cutaneous melanomas.56 The average X-ray attenuation by the iodine and 
average iodine concentration in the lymph nodes with cancer was higher than in 
normal nodes. The presence of cancer within the node did not block uptake of 
the iodinated nanoparticles, as total iodine uptake was higher in cancerous nodes 
with greater than 25% cancer replacement (p < 0.05). The lymph nodes with cancer 
were larger in size, but the uptake of the iodinated contrast agent in the lymph 
nodes was lower. This suggests that the uptake in the region of the cancer was 
lower, but the normal areas became larger and compensated for the portion of 
the lymph node that contained cancer. The architectural alterations in opacified 
cancerous nodes included medullary lymph node filling defects, expansile cortical 
lesions, and disruption of corticomedullary junctions. The authors of this study concluded 
that both quantitative and qualitative differences in iodinated nanoparticle 
enhancement are characteristics that are useful in distinguishing between normal 
and cancerous lymph nodes with CT imaging, following subcutaneous injection of 
iodinated nanoparticles.56 
In an interesting study, iodinated nanoparticles were administered into the 
thorax of dogs using bronchoscopy.59 With this technique, a tube was inserted into 
the lung bronchi under visual guidance with the flexible scope. The nanoparticles 
were injected in the bronchi of the right diaphragmatic lobe of the lung. Two days 
after this procedure, CT images were obtained that did not show any uptake in the 
tracheobronchial lymph nodes. However, images that were obtained after 1 week 
Nanoparticles for Targeting Lymphatics 563 
had excellent contrast enhancement of the tracheobronchial lymph nodes. This contrast 
enhancement remained fixed for 3 weeks following administration.59 The slow 
delivery and prolonged retention of nanoparticles in the tracheobronchial lymph 
nodes has interesting implications for drug delivery. The uptake in the tracheobronchial 
lymph nodes appeared to be secondary to phagocytosis of nanoparticles 
by macrophages. The lymph nodes were enlarged and the histologic specimens 
showed macrophage hyperplasia. Did the macrophages phagocytize the particles 
in the lungs and carry them to the lymph nodes or did the nanoparticles somehow 
get to the lymph nodes, where they became phagocytized? 
5.2. Subcutaneous and intraorgan injection of magnetic resonance 
(MRI) contrast agents 
A limited amount of research has been performed, investigating subcutaneously 
injected gadolinium bound albumin, MS-325, for lymph node diagnosis. In a study 
of normal as well as tumor-bearing hindlegs of rabbits, the subcutaneous administration 
of MS-325 resulted in rapid delineation of popliteal, inguinal, iliac, and 
paraaortic lymph nodes.60 Tumor invasion into lymph nodes presented as magnetic 
resonance imaging (MRI) signal voids in the areas infiltrated by tumor, whereas 
the surrounding residual lymphatic tissue showed enhancement identical to that 
of normal nodes. In addition to providing a safe means of displaying the normal 
lymphatic system, 3D image reconstruction of the MRI image was able to depict 
direct tumor invasion in lymph nodes.60 
Direct injection of magnetic particles into the brain has also been shown to be a 
method for tracking local lymphatic drainage. When small magnetic nanoparticles 
were injected into the brain of rats, up to 50% of the particles were found to drain 
from the CNS via perivascular, perineural and primitive lymphatic drainage to 
the cervical lymph nodes. Central nervous system (CNS) lymphatic drainage may 
occur via connections to the vasculature, but in animal models, up to 50% occurs via 
perivascular, perineural and primitive lymphatic vessels to cervical lymph nodes. 
The trafficking of the superparamagnetic iron particles from the CNS in the rat 
could be visualized both by magnetic resonance imaging (MRI) and histology. These 
magnetic particles appear to provide a tool to rapidly assess drainage of virus-sized 
particles from the CNS using MRI.61 
5.3. Intravenous injection of magnetic nanoparticles for MRI 
imaging 
Ultrasmall nanoparticles can target all of the body's lymph nodes when injected 
intravenously. These ultra-small super-paramagnetic iron oxide particles (USPIO), 
564 Phillips 
also known as ferumoxtran-10, and commercially as Sinerem® in the Netherlands 
(Laboratoire Guerbet, Aulnay sous Bois, France), and as Combidex® in the U.S. 
(Advanced Magnetics, Cambridge, MA), have been developed for improved lymph 
node metastasis detection.62 These particles are composed of 4 to 5 nm iron oxide 
cores surrounded by a dextran coating. After the dextran coating, these particles 
are 20-25 nm in size. These particles produce MRI contrast by shortening the T2 
weighted signal. 
When injected intravenously, these small particles accumulate in the lymph 
node in a high enough concentration to cause significant MRI contrast.63 The percent 
injected dose that ends up in these lymph nodes has not been well studied. In 
spite of lack of quantitation, the uptake in the lymph nodes is sufficient to result 
in effective lymph node contrast in all the lymph nodes of the body. Their half-life 
in circulation is approximately 24 hrs. These intravenously injected nanoparticles 
were associated with a low incidence of adverse reactions. The adverse events most 
frequently seen with USPIO were dyspnea (3.8%), chest pain (2.9%), and rash (2.9%). 
These effects are most likely due to the known complement activation associated 
with intravenous administration of nanoparticles.64 No serious adverse events were 
reported during the 48-hr observation period. There were no clinically significant 
effects on vital signs, physical examination, and laboratory results. Slow infusion of 
a relatively low dose avoids previously reported adverse reactions associated with 
USPIO. 
Nodal accumulation of intravenously injected USPIO is thought to occur, following 
movement of the USPIO through permeable vascular endothelium in lymph 
node vessels of the post-capillary venules and the adjoining capillaries. In one study, 
USPIO were regularly observed at the periphery of the lymph nodes, but not in the 
center of the lymph nodes.65 Isolated iron particles were observed extracellularly 
within lymph vessels in the first hr after injection and by 3 hrs after injection, as 
small dots within macrophages. Numerous dense clusters appeared within the cells 
at later times (i.e. 6 and 12 hrs after injection). These results suggest that the contrast 
agent moves rapidly across the capillary wall to the lymph and is then taken up by 
macrophages.65 
In an initial evaluation of safety and effectiveness of ultrasmall superparamagnetic 
iron oxide particles, 30 adults with suspected lymph node metastasis 
were evaluated with MR imaging before and 22-26 hrs after an intravenous dose 
of USPIO nanoparticles. The sensitivity for metastatic lymph node diagnosis was 
found to be 100% with a specificity of 80%.66 
Another study evaluated USPIO for sensitivity and specificity for differentiating 
metastatic from benign lymph nodes. The study was carried out in 18 
patients with lung cancer. Each patient was evaluated for the homogeneity of the 
lymph node image and change in the post contrast MR signal. All the patients 
Nanoparticles for Targeting Lymphatics 565 
underwent resection of the lymph nodes and histopathologic correlation was performed. 
USPIO was found to have a sensitivity of 92% and a specificity of 80%.62 
Studies have also shown that USPIO are effective in diagnosing metastasis 
in mesorectal lymph nodes. Uniform and central low-signal-intensity patterns in 
lymph nodes are features of nonmalignant nodes. Reactive nodes frequently show 
central low signal with T2-weighted imaging.67 
Recently, a semi-automated technique was developed to detect lymph nodes 
with cancer, following injection of USPIO. Using computer assisted quantitative 
analysis, accurate discrimination between metastatic and normal lymph nodes was 
achieved with a sensitivity of 98% and a specificity of 92%.68 
5.4. Nanopatticle diagnostic agents for localizing the 
sentinel lymph node 
In the last decade, cancer surgeons have become very interested in methods to 
definitively localize the sentinel lymph node. The sentinel lymph node is the first 
lymph node that receives lymphatic drainage from the site of a primary tumor. 
The sentinel node is much more likely to contain metastatic tumor cells than other 
lymph nodes in the same region. It is believed that the initial draining lymph node 
(i.e. sentinel node) of a tumor may reflect the status of the tumor's spread to the 
remaining lymphatic bed. Localization of the sentinel lymph node and its close histological 
assessment, following its removal from the body was initially developed 
as prognostic indicator in patients with malignant melanoma.69 If no cancer cells 
are found in the sentinel node on pathologic examination, the prognosis for the 
patient is greatly improved. After many detailed studies validating the effectiveness 
of this approach for patient prognosis and as a method to guide future therapy 
of melanoma patients, this technique has begun to be applied in other cancers, particularly 
breast cancer. Total lymphadenectomy procedures are being replaced by 
intraoperative lymphatic mapping and sentinel lymph node biopsy.70 
This particular lymph node is now being studied in greater detail, since it is 
able to accurately identify the sentinel node. Close pathological examination of the 
sentinel node with reverse transcriptase-polymerase chain reaction (RT-PCR) has 
shown that the traditional procedures of hematoxylin and eosin (H&E) staining and 
immunohistochemistry underestimate the true incidence of cancer micrometastasis. 
Use of RT-PCR has been shown to be a more powerful predictor of disease 
relapse than traditional H&E staining and immunohistochemical methods.71 The 
advent of the use of sentinel node localization studies has stimulated a general interest 
in lymph node therapy, including both the surgical removal of specific lymph 
nodes as well as lymph node drug delivery. 
566 Phillips 
5.5. Radiolabeled nanoparticles for sentinel lymph node 
identification 
Nanoparticles have played a crucial role in helping to identify and localize the 
sentinel lymph node. This is because many types of nanoparticles have significant 
retention in the first lymph node that they encounter. As much as 40% of the 
nanoparticles that move from the injected site can be retained in the first lymph 
node encountered.16,46 The retention of nanoparticles in the first lymph node is due 
to phagocytosis of the nanoparticles by macrophages.52 The total amount of retention 
in the sentinel node is higher than retention in the subsequent draining lymph 
nodes, because less of the initial dose reaches these secondary nodes. This characteristic 
led to the introduction of radioactive technetium-99m-sulfur colloid particles 
(99mTc-SC) particles as a second method to localize the sentinel lymph node in addition 
to blue dye which was also being used by surgeons.69,72 The use of 99mTc-SC 
in addition to blue dye was a significant improvement over blue dye alone. The 
blue dye was not well retained in the sentinel lymph node and it moved so rapidly 
that the surgeons frequently had difficulty in distinguishing the sentinel node from 
secondary lymph nodes. This problem was solved by the additional use of 99mTc-SC 
which provided a better mark of the sentinel node. Although the blue dye provides 
a desired visual guide for the surgeon, the 99mTc-SC provided verification that the 
correct lymph node had been biopsied. As the 99mTc-SC could be imaged and also 
localized in the operating room prior to and during the surgery, it frequently led 
to a smaller operative incision and a decrease in the time required to find the sentinel 
node.73 The use of radiolabeled nanocoUoids, in addition to blue dye has been 
shown to be complementary techniques that are best used simultaneously.74 
Studies to localize the sentinel lymph node in malignant melanoma and breast 
cancer are now considered the standard of care in patient management. Investigations 
are now being conducted to research the feasibility of using this same 
methodology in many other types of cancer including colon, stomach, head and 
neck, prostate, rectal, lung, uterine, vulvar, and penile cancer. 
5.6. 99mTc-Colloidal nanoparticles for sentinel node identification 
The most common colloidal particle used in the United States is technetium-99msulfur 
colloid (99mTc-SC). 99mTc is readily available, inexpensive and has ideal imaging 
and dosimetry characteristics. Its energy of 140 keV is high enough for emitted 
photons to escape the body without absorption of overlying body tissues, but low 
enough to be readily collimated by lead and absorbed by sodium iodide scintillation 
crystal. Standard SC particles range in size from 10 nm to 1000 nm and they are 
clinically approved for use as liver imaging agents. In this application, the particles 
Nanoparticles for Targeting Lymphatics 567 
are injected intravenously, after which they are rapidly removed by macrophages of 
the liver, spleen and bone marrow. 99mTc-SC makes a physiologic image of the liver 
and prior to the advent of CT scans, it was commonly used to assess the liver for 
tumors. Following the advent of CT scans, it has been used to assess the physiology 
of the liver due to the fact that diseases that damage the liver for any reason, such as 
alcoholic liver disease or infectious hepatitis, also decrease the uptake of nanoparticles 
by the macrophages the are located in the liver, and shift the liver uptake to 
the spleen and the bone marrow. This effect provides a physiologic indicator of the 
health of the liver. 
99mTc-SC has been the nanoparticle used for lymph node identification in the 
United States, because these particles were already widely available and approved 
for clinical use as liver and spleen imaging agents. In the United States, no agent 
has been specifically approved for lymph node detection. These 99mTc-SC particles 
range in size from 10 nm to 1000 nm. However, these particles are not ideal agents 
for the detection of the sentinel lymph node, due to the retention of the majority 
of the injected dose at the peritumoral site of the injection. Studies in animals have 
demonstrated that < 5% of the injected dose (ID) is cleared from the site of injection 
within 60min after injection. With this low clearance, <2% of the injected dose 
accumulates in the sentinel lymph node at 60 min. The intensity of the 99mTc activity 
that is retained at the site of injection frequently makes it hard to locate the sentinel 
lymph node, either by imaging or by the use of a handheld detector probe used 
in surgery. As a result of this problem of retention at the injection site, 99mTc-SC 
particles have been filtered through a 200 nm filter, so that only particles smaller than 
200 nm would be administered.75 These filtered 99mTc-SC nanoparticles have greater 
movement from the injection site, although there is still a debate as to whether 
filtered 99mTc-SC is better than standard 99mTc-SC for sentinel node identification.76 
In Europe, several other colloidal nanoparticles, 99mTc-nanocolloid and 99mTcrhenium 
colloid, are available for use as sentinel node detection agents.77-79 99mTcrhenium 
colloid has an average particle size of less than 100 nm and filtration is not 
required. This 99mTc-rhenium colloid has been shown to be highly effective for the 
identification of the sentinel lymph node.79 
Several other techniques have been investigated that use other nanoparticlebased 
systems to enhance the ability of the surgeon to detect the sentinel node. One 
investigated technique has utilized liposomes containing blue dye to localize the 
sentinel lymph node. Hirnle et al. encapsulated patent blue dye within liposomes 
for potential use in localizing the sentinel lymph node during surgery.80 When this 
technique was studied following injection of the blue liposomes into the lymphatic 
vessels, the lymph nodes were stained blue. Most notably, retroperitoneal lymph 
nodes in rabbits remained dark blue up to 28 days after hindlimb endolymphatic 
instillation of liposomal patent blue dye. This group has also investigated blue 
568 Phillips 
liposome for sentinel node detection in the pig. The blue liposomes were found 
to provide greater intensity blue staining which lasted for a longer duration than 
free unencapsulated blue dye.81 This group later performed a study in humans, 
in which blue liposomes were injected directly into the lymphatic vessels of the 
foot of a patient prior to retroperitoneal staging-lymphadenectomy.82 The lymph 
nodes were well stained with blue dye and were readily visualized at the time 
of the surgery performed 24 hrs following the intralymphatic injection of the blue 
liposomes. 
Plut et al. have also developed a liposome formulation containing blue dye that 
can be radiolabeled with 99mTc.83 The use of the liposome nanoparticle to provide 
a visual identification and the tracking of the liposomes through the lymphatic 
channels, along with the ability to trace the preparation using standard radiation 
detection instrumentation, provides the surgeon with an improved radiolabeled 
compound for lymphoscintigraphy and intraoperative sentinel lymph node identification. 
This method also demonstrates the versatility of nanoparticles to carry 
multiple diagnostic tracers on the same nanoparticle. 
Our group has developed special formulation of liposome encapsulated blue 
dye that can be radiolabeled for imaging and probe detection.44 With this system, 
liposomes trap in the lymph nodes for a prolonged time and they have a high 
efficiency of retention in the lymph nodes. This blue liposome formulation will be 
discussed in detail, in relation to lymph node drug delivery in Sec. 9. 
5.7. Optical 
In a very interesting and recently developed technique, fluorescent quantum dots 
have been investigated as agents to determine the location of the sentinel lymph 
node.84,85 These quantum dots are particles that are 15 nm in diameter. They have 
ideal properties as a non-quenching fluorescent light emitter, following stimulation 
with near infrared light. Animal studies of these quantum dots have shown trapping 
in the sentinel node, following injection in normal animals. They move rapidly to 
the sentinel node so that they can potentially be used during surgery. In one study 
in pigs, the quantum dots were introduced into the lungs as a method of finding the 
sentinel lymph node draining from a lung cancer. The lymph node draining from 
the lung was rapidly identified. This rapid identification of the sentinel lymph node 
using a non-radioactive method was thought to provide a significant advantage for 
the method. 
An important consideration in the future will be the possibility of toxicity from 
the heavy metal, cadmium, which is a major component of these quantum dots. The 
recent synthesis of quantum dots from other agents such as silver may eventually 
lead to quantum dots with improved biological toxicity profiles.86 
Nanoparticles for Targeting Lymphatics 569 
5.8. Ultrasound nanobubbles 
Another type of nanoparticle under investigation for the localization of the sentinel 
lymph node is the nanobubble. Nanobubbles consist of a bilayered shell of 
albumin and an inner layer of a biodegradable polymer known as polycaprolactone. 
This shell encapsulates the gas, nitrogen. Following subcutaneous injection, 
these nanobubbles are tracked using ultrasound imaging. In a study comparing 
microbubbles to sub-micron nanobubbles performed in dogs, the nanobubbles were 
more effective in detecting the sentinel nodes, with 94% or 30 of the 32 sentinel nodes 
being detected. This nanobubble technology could serve as an alternative method 
for detecting the sentinel lymph node. This approach is also unique in that the 
use of ultrasound to detect the bubbles also causes the bubbles to break apart and 
form smaller bubbles.2 It has been proposed that jets released from these nanobubbles 
following exposure to high-frequency ultrasound could be used to nanoinject 
drugs into cells.87 As these nanobubbles can also carry drugs, they could be used 
to deliver high levels of drugs to lymph nodes which could be rapidly released 
following insonation.88 Concerns have been expressed that the energy released 
during insonation of these nanobubbles could produce harmful bioeffects due to 
thermal damage, and further investigation will be required to examine these issues 
in detail.89 
6. Recently Introduced Medical Imaging Devices 
for Monitoring Lymph Node Delivery and 
Therapeutic Response 
For studies of nanoparticle localization of lymph nodes and the assessment of drug 
delivery to lymph nodes for the treatment of cancer metastasis and other lymphatic 
disease processes, the development of clinical imaging systems to direct and verify 
therapy will continue to gain importance. In this regard, the recent advent of clinically 
available single photon emission computed tomography (SPECT)/computed 
X-ray tomography (CT) systems will be important. These systems permit the simultaneous 
image acquisition of a 3D scintigraphic image with a 3D CT image, which 
allows the perfect superimposition of the two types of images. With this new 
SPECT/CT camera, scintigraphic imaging of the location of the nanoparticle can 
be superimposed upon the high resolution CT images which clearly demonstrate 
the anatomy of the body. These systems are already proving to be useful for the 
diagnosis of lymph node metastasis in prostate cancer using targeting radiolabeled 
antibodies.90 The ability to perform a CT image to verify the location of the activity 
visualized on the 3D SPECT scan will be important. Without anatomic verification 
by CT, it is difficult to determine whether the "hot spot" visualized on the image is a 
570 Phillips 
lymph node. It is also difficult to accurately diagnosis the occurrence of a lymphatic 
metastasis. 
By having an imaging machine that can simultaneously co-register a CT image, 
which accurately displays the image, with high resolution and superimposing the 
location of the delivered activity, it will be possible to accurately determine the distribution 
of lymph node directed nanoparticle therapy. In a recent study, the performance 
of SPECT/CT was particularly useful in identifying lymph nodes adjacent 
to the primary lesion. Such nodes are easily overlooked by planar lymphoscintigraphy 
and intraoperative gamma probes, as the high activity at the injection site 
can obscure their detection.91 
Another study also found that SPECT/CT was highly effective for precise preoperative 
localization of the cervical sentinel node in early non-metastatic oral squamous 
cell carcinoma. The authors of this recent study concluded that they believe 
the use of SPECT/CT will become extremely useful, once a consensus has been 
reached on the exclusive excision of the cervical sentinel node in oral cancers, as is 
the case for melanoma or breast cancer.92 
Positron Emission Computed Tomography (PET) has recently entered clinical 
practice for the diagnosis of cancer. Cancer is detected using F-18- 
fluorodeoxyglucose (18F-FDG) due to the greatly increased glucose metabolism of 
cancer cells. This technique is very powerful at detecting cancer in lymph nodes 
containing metastasis.93 It frequently demonstrates cancer in the lymph nodes that 
are not detectable with routine CT imaging. This technique could be very useful in 
monitoring lymph node response, secondary to nanoparticle delivered therapeutic 
agents. Monitoring nanoparticle drug delivery with PET could be particular useful 
in situations where the lymph node if not easily surgically accessible. 
A recent report of a new use of 18F-FDG PET scanning in patients with HIV 
appears to clearly demonstrate specific lymph nodes that are harboring HIV.94 
18F-FDG is a marker of glucose metabolism which is increased in inflammation 
as well as tumors. In this study using PET imaging, the total quantified summed 
activity in lymph nodes was found to correlate with the level of viremia across a 
4-log range. Cervical and axillary lymph nodes were found to have significantly 
more activity than inguinal and ileal nodes (p < 0.0001). The lymph nodes were 
thought to have increased 18F-FDG uptake due to increased glucose metabolism 
in lymph nodes, with lymphocyte turnover secondary to viral infection in these 
infected lymph nodes. This work suggests that the level of virus encountered in the 
lymph nodes is directly related to the blood levels of HIV virus. These PET studies 
also clearly indicated that particular lymph nodes were involved, suggesting 
the possibility that these nodes could be specifically targeted with nanoparticles 
for drug delivery. A second article investigating PET scanning in patients with HIV 
has even suggested that surgical intervention to remove the specific nodes should be 
considered. An alternative strategy would be to use nanoparticles to target antiviral 
Nanoparticles for Targeting Lymphatics 571 
medications to specific lymph nodes, based on the PET imaging data. Imaging with 
SPECT/CT could be used to determine that the nanoparticle associated drugs were 
specifically delivered to the targeted lymph node, and repeat PET imaging could 
be used to monitor the effectiveness of the delivered therapy for the treatment of 
the lymph node. 
Analogues of these human imaging systems have been developed for small 
animal imaging research. These commercially available imaging systems combine 
microSPECT and microCT into the same unit for image superimposition. 
A new commercially available imaging system has become available, combining 
microSPECT, microCT and microPET into the same unit. These small animal imaging 
systems will prove to be valuable in the investigation of nanoparticle lymph 
node delivery. They should facilitate the translation of basic research to imaging 
studies in the clinical environment. 
7. Nanoparticle Lymph Node Drug Delivery 
Several authors have recommended that in the case of diseases with lymphatic 
involvement, it is desirable to develop treatment approaches to deliver diagnostic 
and therapeutic agents to lymph nodes that could help prevent diseases from 
spreading.8,10,13'17 This situation has motivated interest in developing methods for 
specific drug delivery to lymph nodes and lymphatic vessels, with the hope of 
improving cancer and infection control and treatment. 
Whatever the carrier system chosen, the following basic requirements need 
to be met in order to develop an effective drug delivery system: (a) the carrier 
system must concentrate selectively in the target organ or tissue; (b) the carrier 
system must release the drug in such a way as to achieve its chemotherapeutic or 
pharmacological effect, over the required time frame; (c) the carrier system must be 
stable before administration and during its journey to the target, but able to release 
the drug when it is at the target; and (d) the carrier must also be compatible with 
the body in terms of toxicity, biodegradation and immunogenicity.11 Nanoparticles 
of different types can meet the above criteria as effective carriers for drug delivery 
to lymph nodes. Targeting of lymph nodes for drug delivery has been attempted 
by the use of emulsions, non-lipid macromolecules, antibodies, and nanoparticles. 
Nanoparticles appear to hold the most promise for lymphatic drug delivery, in 
comparison with other types of carriers. In the present review, we will focus on the 
variety of nanoparticle systems that have been used for lymph node drug delivery. 
7.1. Confusion in reporting lymph node delivery 
The research literature reporting nanoparticle uptake in lymph nodes can be very 
confusing. This is due to the tendency of many investigators to report the percent 
572 Phillips 
uptake in the lymph node as a percent of the injected dose per gram of tissue. 
Although reporting uptake as a percent of the injected dose per gram is sensible 
from a tissue treatment standpoint, it does not easily allow the readers to determine 
the percent of the total administered dose accumulated in the lymph node. For 
instance, in animals such as mice with very small lymph nodes that weigh only a 
fraction of a gram, the accumulated doses can be as high as 100% of the injected dose 
per gram, but considering that a mouse lymph node weighs only 0.01 gram this is 
only 1% of the total dose that was administered (1% ID). Had this same fraction 
of the injected dose accumulated in a rat lymph node that weighs approximately 
0.1 gram, this dose would have been only 10% of the injected dose per gram; and 
in a human with a lymph node with a weight of 1 gram, it would have been only 
1% ID per gram. It is very important to keep these species differences in mind 
when interpreting the previous literature and it would be best if all investigators 
would report their research not only in terms of dose per gram, but also in terms 
of the percentage of the total injected dose delivered to the lymph node. From a 
pharmacologic standpoint, percent dose per gram may be considered correct, but it 
is highly unlikely that these percent dose per gram results in mice would translate 
into humans. Based on our experience, it is much more likely that the percent that 
clears from the injected site will always be approximately the same in each species 
and the percent that accumulates in the lymph nodes, no matter what its weight in 
grams, will also be approximately the same. 
For example, in one study, researchers reported that subcutaneously injected 
liposomes without a lymph node targeting mechanism had much higher concentration 
in the lymph nodes on the side of the subcutaneous injection, compared with 
the lymph nodes of the opposite side that did not receive the injection.95 24 hrs after 
subcutaneous injection, 57.9% of the ID/gram of tissue was found in the inguinal 
node on the side of the injection (ipsilateral side) versus only 0.48% ID/gram of 
tissue in the lymph node of the opposite side of the injection (contralateral side) at 
24 hrs. Again, this study was carried out in mice and the % ID per gram is somewhat 
deceptive due to the fact that mice have very small lymph nodes. This probably 
represents no more than 1-3% of the total injected dose accumulating in the lymph 
node that drained from the subcutaneous site. The important point is unchanged. 
There was more than a 100-fold increased amount of liposomes deposited on the 
side of subcutaneous injection, compared with the lymph node on the other side 
that did not receive the subcutaneous injection.95 
It is important that readers of the lymph node drug delivery literature take 
lymph node weight into consideration. For instance, if 20% of the total subcutaneous 
injected dose were to accumulate in the lymph node of a mouse with a lymph node 
that weighs just 0.01 gram, the % ID/gram would be 20 divided by 0.01 or 2000% 
of the injected dose per gram. 
Nanoparticles for Targeting Lymphatics 573 
7.2. Calculation of lymph node retention efficiency 
Using scintigraphic imaging, our group has developed a method to calculate lymph 
node retention efficiency.3 The calculation describes the fraction of the nanoparticles 
that are cleared from the initial subcutaneous site of injection that become trapped 
and retained in the lymph node. This estimated lymph node retention calculation 
describes how efficiently the lymph node can retain a particular subcutaneously 
injected nanoparticle. It also describes the portion of the dose that enters a lymph 
node and then leaves that primary lymph node and moves to the next lymph node. 
The calculation requires that the nanoparticle be labeled with a radiotracer and 
imaged scintigraphically. It is determined by drawing a region of activity around the 
injection site and determining the total percentage of the baseline injected activity 
that has cleared the injection site at various times post injection. This total cleared 
activity is then assumed to be the amount that enters the first lymph node. The 
activity retained in the lymph node is divided by the total activity cleared from 
the injection site, so that a lymph node retention efficiency can be calculated. With 
most unmodified liposome preparations, lymph nodes only retain approximately 
4% of the liposomes that enter the lymph node. Retention efficiency is higher with 
other particles, such as unfiltered 99mTc-SC that have a 40% lymph node retention 
efficiency, but these particles are also associated with a very poor clearance from 
the initial injection site.16 
8. Specific Types Nanoparticles for Lymph Node Targeting 
8.1. PLGA nanoparticles 
Poly(lactide-co-glycolide)(PLGA) nanoparticles have been investigated for lymph 
node drug delivery agents. By modifying the surface of the PLGA nanoparticles 
with block co-polymers, PLGA nanoparticles that deliver up to 17% of a subcutaneously 
injected dose to regional lymph nodes have been developed.96 Lymphatic 
uptake was studied by labeling these PLGA nanoparticles with mIn-oxime. Lymphatic 
uptake of all coated PLGA nanoparticles was enhanced, compared with the 
uncoated PLGA nanoparticles. This research suggests that the nanospheres are suitable 
for diagnostic and therapeutic lymph node delivery applications in clinical and 
experimental medicine.96 
In another study, the in vivo trafficking of PLGA particles, encapsulating a diphtheria 
antigen, was possible using a fluorescent marker following subcutaneous 
administration for the immunization of mice,97 these fluorescent PLGA microspheres 
were observed in cells of lymphoid tissues such as mesenteric lymph nodes 
and spleen. However, particle fluorescence in lymphoid tissues decreased rapidly, 
as they were degraded inside the cells, thereby enabling the presentation of the 
574 Phillips 
antigen to specific cells of the immune system. This is one of the few studies in 
which a nanoparticle carrier for vaccine delivery was tracked in vivo.97 
8.2. Micelles 
Particles known as micelles can be formed from amphiphilic biocompatible polyoxyethylene 
(PEO)-based polymers. These nanoparticles (micelles) are 10-50 nm 
in diameter. Researchers have labeled micelles with amphiphilic indium-Ill (mIn) 
and gadolinium chelates and used them as nanoparticulate contrast media for imaging 
lymph nodes, following subcutaneous injection into the rabbit's paw.98 Corresponding 
images of local lymphatics were acquired using a gamma camera and a 
magnetic resonance (MR) imager. The entire lymphatic vessel chain from the paw 
to the thoracic duct could be visualized using m I n labeled micelles. After injection 
site massage, 40% of the injected dose cleared from the injection site into the lymphatics, 
30 min after massage was performed over the site of injection. Although 
approximately 8% of the subcutaneously injected dose was in the popliteal lymph 
node at 30 min, this amount decreased significantly after massage over the popliteal 
lymph node, with less than 2% of micelles remaining in the lymph node. When 
gadolinium containing micelles were administered, Tl-weighted MR images of the 
primary lymph node had significant contrast enhancement within 4 min following 
massage.98 Prior to this imaging research, Torchilin had previously shown that 
micelles have much promise for the delivery of poorly soluble drugs following 
intravenous delivery.99 Micelles also appear to be promising agents for lymphatic 
drug delivery, particularly if methods can be developed to increase their retention 
in lymph nodes. 
8.3. Liposomes 
The investigation of liposomes as carriers for lymph node drug delivery was first 
performed by Segal et a/.100 Following the intratesticular injection of liposomes 
encapsulating the anti-cancer drug actinomycin D, high concentrations of the drug 
were found in the local lymph nodes. Lymph node imaging using 99mTc-labeled 
liposomes was first performed by Osborne et a/.101 Liposome distributions were 
determined in rats following injection of the 99mTc liposomes in the rat hind footpads. 
In these studies, 1 to 2% of the injected dose of neutral and cationic liposomes 
were found to localize in the draining lymph nodes. Negatively charged liposomes 
did not show good accumulation in the lymph nodes.101 Soon after these studies 
were performed, the reliability of these previous studies for representing the actual 
distribution of liposomes was questioned and it was suggested that much of the 
activity localized in the lymph nodes was not associated with liposomes, due to 
Nanoparticles for Targeting Lymphatics 575 
instability of the 99mTc label.102 This article recommended that new methods of 
labeling liposomes with 99mTc be developed. Follow-up studies demonstrated that 
the type of labeling used in the prior studies, in which 99mTc was labeled to the outer 
surface of liposomes following reduction of the 99mTc with stannous chloride, was 
not stable.103 It was particularly unstable for labeling liposomes with a cationic and 
neutral surface charge, which possibly explained the low accumulation in lymph 
nodes found with anionic liposomes. These studies demonstrate the importance of 
label stability in the tracking of liposomes for quantitation of targeting to the lymph 
nodes. 
Since those early studies, more stable methods of labeling liposomes with 
99mTc have been developed and applied to lymph node imaging. A method developed 
by our group uses hexamethylpropyleneamine oxime (HMPAO), a clinically 
approved and commercially available chelator of 99mTc used for brain imaging.104 
In this method, 99mTc-pertechnetate, which is readily available from a generator, is 
incubated for 5 min with HMPAO, which chelates the 99mTc into lipophilic 99mTc- 
HMPAO.105 The 99mTc-HMPAO is then added to previously manufactured liposomes 
that encapsulate glutathione. It is generally believed that lipophilic HMPAO 
carries the 99mTc across the lipid bilayer of liposome into the aqueous interior of 
the liposome where it interacts with the encapsulated glutathione, resulting in its 
conversion to hydrophilic 99mTc-HMPAO. The hydrophilic 99mTc-HMPAO is irreversibly 
trapped in the aqueous phase of the liposome because it is unable to cross 
the lipid membrane. A similar mechanism has been proposed to explain the process 
whereby 99mTc-HMPAO becomes trapped in brain cells for use as a brain imaging 
agent. This liposome label is very stable with minimal dissociation of the 99mTc 
from the liposomes. It has been used to study the distribution of intravenously 
administered liposomes.106 
Using the above liposome labeling methodology, detailed studies were carried 
out to assess the effect of liposome size and surface modifications on movement 
from the subcutaneous site of injection, as well as the retention of the liposomes in 
the lymph node.16 The fraction of liposomes that are cleared from the subcutaneous 
injection site depends on the size and surface characteristics of a particular liposome 
formulation. Their large size of liposomes (> 80 nm) precludes their direct absorption 
into the blood stream. In studying a wide range of liposome sizes from 86 nm 
liposomes to 520 nm liposome, there was little difference in the ultimate accumulation 
of liposomes in the first or sentinel lymph node at 24 hrs post administration. 
This lymph node retention ranged from 1.3% to 2.4% of the total administered 
dose.16 This retention in the lymph node is relatively low, considering that most 
subcutaneously injected liposome preparations more than 50% of the injected liposomes 
are cleared from the injection site. The lymph node retention for unmodified 
liposomes is low, compared with other colloidal particles such as 99mTc-SC. There 
576 Phillips 
were differences between the different liposomes sizes and surface characteristics 
in regard to the percent of activity that cleared from the subcutaneous site of injection. 
Small liposomes and those coated with PEG had the greatest clearance from 
the subcutaneous site of injection, with small 86 nm liposomes having < 40% of the 
injected dose remaining at the injection site at 24 hrs. Larger neutral and negatively 
charged liposomes had > 60% of the injected dose remaining at the initial site of 
subcutaneous injection. However, this smaller amount of large liposomes that were 
cleared from the injection site was compensated for by better retention in the lymph 
node.16 
It appears that the properties of liposomes which enhance their clearance from 
the injection site, also decrease their retention in the lymph nodes. The generally 
low overall lymph node retention of most standard liposome formulations is likely 
to be due to their natural lipid composition, which probably allows a large percentage 
of liposomes that enter a lymph node to escape recognition and phagocytosis 
by macrophages that line the endothelium of the lymph node. This relatively 
low retention of liposomes in lymph nodes has also been reported by Ousseren 
et al.5-i5'A7'W7 
One of the several liposome surface modifications that have resulted in modestly 
increased retention of subcutaneously injected liposomes in the lymph node 
is the use of positively charged lipids in the liposome. Liposomes containing positively 
charge lipids had approximately 2 to 3 times the lymph node localization (up 
to 3.6% of the injected dose), as liposomes containing neutral or negatively charged 
lipids (1.2% of the injected dose).47 Coating liposomes with the antibody, IgG, has 
been shown to increase the lymph node localization of liposomes to 4.5% of the 
injected dose at 1 hr, but this level decreased to 3% by 24 hrs.108 Attaching mannose 
to the surface of a liposome has also been reported to modestly increase lymph 
node uptake by 3-fold, compared with control liposomes.109 None of these previously 
mentioned modifications has resulted in large increases in the percentage of 
liposomes deposited in the draining lymph nodes, while most of the lymphatically 
absorbed liposome dose passes through the lymph nodes. 
This relatively low retention may be due to the fact that the natural lipid composition 
of liposomes allows them to escape recognition by the macrophages located 
in the lymph nodes. In comparison with other particles, an abundant fraction of 
the subcutaneously injected dose of liposomes moves into the lymphatic system, 
so that greater than 50% clears from the site of injection by 24 hrs. Even though this 
retention of liposomes is very low in comparison with the injected dose, its uptake 
can still be quite substantial in terms of drug delivery, when considered on a per 
gram of tissue basis. For example, 1 to 2% of the injection dose in each lymph node 
represents a total per gram tissue uptake that is generally greater by 30-40 fold than 
Nanoparticles for Targeting Lymphatics 577 
the liposome uptake that will eventually reach the liver and spleen following the 
same subcutaneous injection. 
9. Avidin Biotin-Liposome Lymph Node Targeting Method 
The relatively low retention of liposomes in lymph nodes led our group to search 
for new ways to improve liposome retention in lymph nodes. This research resulted 
in development of a new method of increasing liposome retention in lymph nodes 
following subcutaneous injection.3 This lymph node targeting method utilizes the 
high affinity ligands, biotin and avidin. Biotin is a naturally occurring cofactor and 
avidin is a protein derived from eggs. Avidin and biotin have an extremely high 
affinity for each other. Avidin has 4 receptor sites for biotin associated with each 
molecule. These 4 receptor sites permit the binding of multiple biotin molecules 
which causes aggregation of liposomes that have biotin on their surface. Using 
this method, liposomes coated with biotin on their surface (biotin-liposomes) are 
injected subcutaneously, followed by a nearby subcutaneous injection of avidin. Following 
their subcutaneous injection, the avidin and the biotin-liposomes move into 
the lymphatic vessels. It was orginally hypothesized that the biotin-liposomes that 
are migrating through the lymphatic vessels meet with the avidin, resulting in an 
aggregate that becomes trapped in the lymph nodes. Subsequent research suggests 
that an alternative possibility maybe more likely.110,111 This alternative hypothesis is 
that the positively charged avidin becomes bound to negatively charged endothelial 
cells in the lymph nodes and the biotin-liposomes become bound by these avidin 
molecules attached to the endothelial surface. It is possible that both processes are 
occurring, however, research with intracavitary avidin/biotin-liposome systems 
suggests that the second possibility may be more likely.110,111 
This in vivo nanoassembly of biotin-liposome/avidin aggregates mimics processes 
that occur naturally in the body such as the aggregation of platelets and the 
aggregation of infectious agents by antibodies. The biotin-liposome/avidin system 
has promising potential for therapeutic agent delivery to lymph nodes. It can be 
applied not only to subcutaneous targeting of lymph nodes, but also to intracavitary 
lymph node targeting. Although liposomes are the particular nanoparticle in 
which this methodology has been developed, it should also be possible to apply 
this methodology to the other types of nanoparticle carriers. Other high affinity 
ligands pairs as alternatives to avidin and biotin could also be used. Scintigraphic 
imaging of liposomes labeled with 99mTc, labeled in a stable fashion, has greatly 
aided the determination of the proper concentration of avidin and biotin and could 
be used to develop similar targeting methodology with other nanocarriers.3 
578 Phillips 
As an extension of the avidin/biotin-liposome lymph node targeting system, 
we have developed a special liposome formulation that contains both encapsulated 
blue dye as a potential system for localizing the sentinel lymph node, visually as 
well as scintigraphically, and/or with a gamma probe.44 Potential advantages of 
this system over the current methods are that it can be performed any time from 
1 hr to 1 day before the surgery is planned, because the lymph nodes are stained 
blue for a prolonged time and the sentinel lymph has the highest concentration of 
liposomes. Using this method, a separate blue dye injection just prior to surgery 
would not be necessary. 
Examination of blue lymph nodes by light microscopy reveals that the liposomes 
tend to be deposited only in the outer cortex of the lymph node, however, 
this is not always the case as lymph nodes can be completely stained, depending 
on the concentration and timing of the avidin and biotin-liposome injection. 
As part of these studies, we have found that lymph nodes remain blue by visual 
observation in vivo, for more than a week following subcutaneous injection. The 
prolonged retention and slow release observed with blue biotin-liposomes demonstrates 
the potential of this system for the delivery and sustained release of drugs 
in the lymph nodes. Clinical studies would be required to determine whether the 
biotin-liposome/ avidin system is effective in targeting the sentinel node in humans. 
10. Massage and the Avidin-Biotin Liposome 
Targeting Method 
Repeated massage and reinfusion of saline in a subcutaneous site of injection of the 
biotin-liposomes and avidin can be used to further enhance lymph node accumulation, 
as well as the rapidity at which the liposomes leave the subcutaneous site. In 
a study in our laboratory, the effect of this repeated massage on lymph node accumulation 
with the avidin-biotin-liposomes technique was compared with the same 
method with filtered 99mTc-SC particles. In this study, a 24-gauge angiocatheter was 
placed subcutaneously in the dorsal foot of an anesthetized rabbit and 0.3 ml 99mTcradiolabeled 
biotin-liposomes or 99mTc-SC were infused through the catheter into 
the subcutaneous tissue. Avidin (5 mg in 0.3 ml) was injected into the rabbit hind 
foot, approximately 2 cm proximal to the biotin-liposome injection site, in the same 
manner as described in the initial biotin-liposome study.3 The injection site was 
massaged for 5 min. At serial 5 min intervals, 0.3 ml of saline was infused through 
the indwelling subcutaneous catheter and massage was performed for 3 min, following 
each reinfusion of saline for the first hr. During the second hr post injection, 
this 0.3 ml saline infusion followed by massage was repeated at 10-min intervals 
until 120 min post injection. Images were acquired at baseline, 30, 60 and 120 min 
Nanoparticles for Targeting Lymphatics 579 
and at 24 hrs. After acquiring images at 2 hrs, the rabbits recovered from the anesthesia 
and were returned to the animal facility until they were imaged at 24 hrs. 
As shown in Fig. 3, the massage more than doubled the liposome accumulation 
in the popliteal node of the rabbits. The intensity of 99mTc activity in the popliteal 
node is equal or greater than that of the activity at the injection site. At 2 hrs, an 
average of 29% of the injected dose of liposomes were trapped in the popliteal node 
of the foot that received the serial massages and saline infusions versus 13% in the 
popliteal node on the side that was not massaged. At 2 hrs, more than 2.5 times 
the injected dose had cleared from the site of injection in the foot on the side with 
the serial saline infusions and massage, compared with the control side (50% versus 
20%), so that massage with serial saline infusions can greatly increase the rate and 
total movement from the injection site. The liposome uptake in the popliteal node 
was generally fixed for an extended period. At 24 hrs, 20% of the injected dose was 
still retained in the lymph node that received the repeated massage and saline infusions. 
When this same methodology was compared in filtered 99mTc-SC particles, 
the greatly increased accumulation in the lymph nodes was not observed. Although 
a greater amount of filtered 99mTc-SC particles was cleared from the subcutaneous 
injection site in the foot with the serial saline infusions and massage, (37% vs 20%), 
2 Hours Post 
Subcutaneous Injections in Feet 
Biotin Liposome Sulfur Colloid 
h-Liposome Inj. FeeH 
No Saline + No Saline + 
Massage Massage Massage Massage 
Fig. 3. Scintigraphic images of rabbit feet and legs acquired 2 hrs after administration of 
99mTc-biotin liposomes subcutaneously in the feet of a rabbit demonstrating the effect of 
repeated massage and saline infusion on lymph node accumulation following injection 
of radiolabeled biotin liposomes and avidin was compared with the effect of massage on 
filtered 99mTc-SC particles. The massage and repeat saline infusions more than doubled the 
liposome accumulation in the popliteal node of the rabbits (29% vs 11%). Massage and saline 
infusion did not greatly increase the lymph node accumulation of filtered 99mTc-SC particles 
(13% vs 9%). It appears that the extra saline and massage simply pushed most of the 99mTc-SC 
through the lymph node. This was unlike what happened with the biotin-liposomes, where 
the avidin continued to trap or aggregate the biotin-liposomes in the lymph node. 
580 Phillips 
the percent of the injected dose retained at 2 hrs in the popliteal node was only 
minimally increased (13.2% vs 9%). It appears that the extra saline and massage 
simply pushed most of the 99mTc-SC through the lymph node. This was very different 
from the biotin-liposomes, where the avidin continued to trap or aggregate 
the biotin-liposomes in the lymph node. 
11. Nanoparticles for Lymph Node Anti-Infectious 
Agent Delivery 
Only a few studies have examined the delivery of nanoparticles to lymph nodes for 
the treatment of infectious disease. In one study, liposome encapsulated amikacin 
was injected subcutaneously, intramuscularly, and intravenously. Drug levels in the 
lymph nodes were studied at various time points following injection. Drug level 
area under the curve (AUCs) in regional lymph nodes exceeded plasma AUCs by 
4-fold, after subcutaneous and intramuscular injection of liposomal amikacin.112 
The authors of the study conclude that liposomes encapsulating amikacin have 
much potential for drug delivery, and even suggest that these liposomes could 
potentially be used for local delivery in perioperative prophylaxis, pneumonias 
and intralesional therapy as well as sustained systemic delivery of encapsulated 
drugs.112 
This effectiveness of liposome encapsulated amikacin following subcutaneous 
injection differs significantly from a previous study, in which 400 nm liposome 
encapsulated amikacin was administered intravenously. The intravenously administered 
amikacin encapsulated liposomes were effective against M. avium intracellulare 
located in the liver and spleen, but they had no effect on the organisms that 
were located in the lymph nodes.113 It is unlikely that these intravenously injected 
liposomes accumulated in the lymph nodes to any degree. 
The use of nanoparticles to increase the drug delivery to HIV infected lymph 
nodes appears promising. When injected intravenously, the anti-HIV drug, indinavir, 
was found to achieve drugs levels in lymph node mononuclear cells that 
were only 25-35% of mononuclear cells in the blood. Lipid suspensions of the antiviral 
HIV drug, indinavir, form particles that are 50-80 nm in size. When these 
particles were injected subcutaneously in HIV infected macaques, very high drug 
levels were achieved in the lymph nodes and viral RNA loads in these nodes were 
greatly reduced. This lowering of HIV viral RNA levels could not be achieved when 
the same drug was injected intravenously.114 
Dufresne et al. have investigated liposomes coated with anti-HLA-DR Fab' 
fragments for specifically targeting liposomes to follicular dendritic cells and 
macrophages within the lymph nodes of mice, with the goal of increasing the delivery 
of antiviral drugs to these cells infected with HIV115 The uptake of anti-HLA-DR 
Nanoparticles for Targeting Lymphatics 581 
Fab' coated liposomes within lymph nodes was 2 to 3-fold higher when compared 
with conventional liposomes. However, of more importance is the potential specific 
delivery of the anti-HLA-DR Fab' liposomes to antigen presenting cells within the 
lymph node. 
More recently, researchers from this same group have investigated the targeting 
of lymph nodes with indinavir, a protease inhibitor, encapsulated into immunoliposomes 
coated with the same anti-HLA-DR Fab' antibody fragment. Mice were 
injected subcutaneously below the neck with either free indinavir or liposome 
encapsulated indinavir. Animals were sacrificed at various times following injection 
and tissues collected and analyzed for indinavir drug levels. Drug levels were 
compared in lymph nodes from the mice receiving the subcutaneously injected free 
drug and subcutaneously injected liposome encapsulated drug. Drug levels in the 
brachial and cervical lymph nodes were 126 and 69 times greater with the liposome 
encapsulated drug, than the free drug.24 
12. Liposomes for Intraperitoneal Lymph Node Drug Delivery 
Intraperitoneal drug delivery is currently considered a viable approach for the treatment 
of ovarian cancer.116-118 Studies in which free drugs are administered into the 
peritoneum have shown survival benefits in ovarian cancer patients. Although 
most intraperitoneally delivered unencapsulated free drugs are rapidly cleared 
from the peritoneal fluid without entering the lymphatic system, direct intraperitoneal 
administration of drugs can achieve much higher peak concentrations in 
the peritoneal fluid, compared with the same drug administered intravenously 
(20-fold higher for cisplatin and carboplatin, to as high as 1000-fold for taxol).116-118 
Although these drug levels quickly equilibrate with plasma after termination of the 
peritoneal infusion, transiently elevated peritoneal drug levels provide a significant 
therapeutic advantage. These have led many investigators to be enthusiastic about 
this approach for ovarian cancer treatment.117119 Unfortunately, rapid clearance of 
these free drugs from the peritoneum diminishes the advantages derived from the 
intraperitoneal infusion procedure. One way of improving peritoneal drug delivery 
could be through the use of nanoparticle drug carriers.120-122 
Nanoparticles have significant promise as carriers for intraperitoneal drug 
delivery. Many disease processes spread by dissemination through the peritoneum. 
For instance, dissemination of cancer cells throughout the peritoneum is a very 
common manifestation of ovarian and gastric cancer.123 When the cancer cells 
spread throughout the peritoneum, they are frequently trapped in lymph nodes that 
receive peritoneal fluid drainage.124 The normal pathway of drug clearance from 
the peritoneum is either through direct absorption across the peritoneal membrane 
or by drainage into the lymphatic system through absorption by the stomata in the 
582 Phillips 
diaphragm.125 These diaphragmatic stomata are fairly large and can absorbed red 
blood cells from the peritoneal fluid.126'127 
Intraperitoneally administered therapeutic nanoparticles not only increase and 
prolong the drug delivery in the peritoneum, but they can also increase delivery 
of therapeutic agents to the lymph nodes that filter lymph fluid drainage from the 
peritoneum. These lymph nodes frequently contain cancer cells. 
12.1. Intraperitoneal liposome encapsulated drugs 
Intraperitoneal administration of a liposome encapsulated drug not only increases 
the retention of the drug in the peritoneum, it also increases the delivery of the 
drug to lymph nodes that drain from the peritoneum. This is because liposome 
encapsulated drugs are mostly cleared through the lymphatic vessels, with at least 
a portion of the administered drug being deposited in the lymph nodes, where it 
degrades and is slowly released from the liposome and lymph node macrophages 
in high concentrations.121,128 
Other studies demonstrate an improved toxicity profile. For instance, encapsulation 
of paclitaxel in a liposome has been shown to have decreased toxicity 
following intraperitoneal administration, while retaining equal efficacy for the 
treatment of intraperitoneal P388 leukemia.129 In humans, the dose limiting toxicity 
from intraperitoneal administration of paclitaxel was severe abdominal pain, 
which was thought to be due to direct toxicity from either the paclitaxel or the 
ethanol/polyethoxylated castor oil delivery vehicle.130 
Intraperitoneal delivery has also shown promise for nanoparticle gene transfection 
with novel cationic lipid containing liposomes.131 These cationic liposome 
contained luciferase and beta-galactosidase genes that served as reporter 
genes. Intraperitoneal gene delivery for peritoneal disseminated ovarian cancer in 
nude mice was achieved using a stable chloramphenicol acetyl transferase (CAT)- 
expressing ovarian cancer cell line (OV-CA-2774/CAT), which permitted quantification 
of the exact tumor burden of organs. Intraperitoneal gene delivery to these 
disseminated ovarian cancer cells was excellent, with gene transfection appearing 
to be specific to intraperitoneal ovarian cancer cells. The 0-Chol:DNA lipoplex 
appears to offer potential advantages over other commercial transfection reagents 
because of (1) its higher level of gene expression in vitro and in vivo; (2) its reduced 
susceptibility to serum inhibition; and (3) its highly selective transfection into tumor 
cells. These results suggest that the 0-Chol:DNA lipoplex is a promising tool in gene 
therapy for patients with peritoneal disseminated ovarian cancer.131 
An important potential application of the intraperitoneal delivery of liposomes 
and other nanoparticles that carry anti cancer agents is in the prophylaxis 
of peritoneal carcinomatosis. As 50% of patients with malignant gastrointestinal 
Nanoparticles for Targeting Lymphatics 583 
or gynecological diseases experience peritoneal carcinomatosis shortly after local 
curative resection, there is a great interest in delivering intraperitoneal chemotherapy 
during the perioperative period. One study found that the intraperitoneal 
administration of the chemotherapeutic agents, cisplatin and mitomycin, prevented 
perioperative peritoneal carcinomatosis in a rat model.132'133 The rats receiving 
cisplatin did, however, experience severe, local toxicity with bleeding into the 
peritoneum and toxic necrotic reactions of the colon. Liposomes encapsulating 
anticancer agents could potentially be used for this type of perioperative chemotherapy. 
The potential for the treatment of micrometastasis in lymph nodes with this 
liposomes is also great. Metastasis to mediastinal and other lymph nodes receiving 
lymph drainage from the peritoneal fluid are not uncommon findings in ovarian 
cancer at autopsy.124 
One important consideration that might influence the effectiveness of intraperitoneal 
lymph node drug delivery is the possibility that the lymphatics could be completely 
obstructed with tumor, and therefore not accessible for lymph drainage. 
Generally, the ascites that develops in patients with intraperitoneal cancer are 
thought to be due to the obstruction of the lymphatics by the metastatic cancer. 
In a recent study, only 12.5% of women diagnosed with early stage ovarian cancer 
presented with ascites.134 Biodistribution studies with liposome imaging could be 
performed to determine the effectiveness of lymph node targeting in situations of 
suspected lymph node obstruction. 
12.2. Effect of liposome size on intraperitoneal clearance 
It must be noted that simply making liposomes larger does not increase retention in 
the peritoneum or lymph nodes that receive drainage from the peritoneum. Hirono 
and Hunt have performed an extensive study on the effect of liposome size on 
their subsequent distribution, after intraperitoneal administration of liposomes of 
different sizes. In their studies, 50-60% of the intraperitoneal dose of liposomes 
of varying sizes encapsulating carbon-14 (14C) labeled-sucrose cleared from the 
peritoneum by 5 hr in all liposomes studied. These liposomes ranged in size from 
48 nm to 720 nm. The greatest amount of 14C-sucrose (~ 40%) appeared in the urine 
after administration of the largest liposomes. The authors speculated that the large 
460 nm and 720 nm liposomes were unstable in the peritoneum, so that they rapidly 
released their encapsulated 14C-sucrose. It is also unlikely that simply increasing 
the size of the liposomes, in and of itself, would be sufficient to result in increased 
peritoneal and lymph node retention, because particles as large as erythrocytes 
have been demonstrated to readily drain from the peritoneum, by passing through 
the diaphragmatic stomata and returning to the blood stream. In one study of 
584 Phillips 
chromium-51 labeled red blood cells injected into the peritoneal cavity of sheep, 
80% of the red cells returned to the blood circulation by 6 hrs after administration.126 
12.3. Avidin/Biotin-liposome system for intraperitoneal and lymph 
node drug delivery 
Few of the above previously described studies with intraperitoneally administered 
liposome nanoparticles have focused on the fact that liposome nanoparticles clear 
from the peritoneum by passing through the lymphatic vessels. The liposomes pass 
through and are only partially trapped to lesser or greater degrees in the lymph 
nodes that drain from the peritoneum. These lymph nodes frequently contain cancer 
metastasis. 
Our group has applied the previously described avidin/biotin-liposome system 
to intraperitoneal drug delivery.18 The intraperitoneal biotin-liposome/avidin 
delivery method previously described in this paper has potential as a delivery system 
for the local treatment of intraperitoneal and intralymphatic disease processes, 
by increasing the retention of drugs in the peritoneum and in the lymph nodes that 
receive lymphatic drainage from the peritoneum. 
The interaction of biotin-liposomes with avidin apparently results in the aggregation 
of the liposomes in the peritoneum. This aggregation greatly alters the distribution 
of liposomes and appears to result in a prolonged retention of liposomes in 
the peritoneum, as well as an increased accumulation and retention of liposomes in 
lymph nodes receiving drainage from the peritoneum. Rats that received intraperitoneal 
injection of biotin-liposomes and avidin had only a minimal percentage of 
the injected dose of liposomes that reach the systemic circulation by 24 hrs and a 
low % ID was found in the spleen, blood and liver at 24 hrs (< 9% ID). In contrast, 
control animals, administered only the biotin-liposomes without the avidin, had 
23% ID in the spleen, 14% ID in the blood, and 9.8% ID in the liver, for a total of 47% 
ID in these organs at 24 hrs. Lymph nodes in the abdomen and in the mediastinum 
in rats receiving avidin had also greatly increased uptake of the biotin-liposomes. 
Delivery of liposome encapsulated drugs using this method should provide 
sustained local release of drug within the peritoneum and the lymph nodes 
draining the peritoneum, as the liposomes degrade or become phagocytized by 
macrophages. This delivery system could also attenuate systemic drug toxicities 
by greatly reducing the rate at which a drug returns to the systemic circulation by 
either preventing rapid direct absorption through the peritoneal membrane, or by 
moderately preventing rapid passage through the lymphatic vessels and lymph 
nodes back to the blood. 
Intraperitoneal administration using the avidin/biotin-liposome system has 
potential as a carrier system for the delivery of anti cancer agents in the peritoneum, 
Nanopa nicies for Ta rgeting Lympha tics 585 
as well as liposome encapsulated radiotherapeutic beta-emitters for the treatment 
of peritoneally disseminated ovarian and upper gastrointestinal cancers. 
12.4. Mediastinal lymph node drug delivery with avidin-biotin 
system by intrapleural injection 
Mediastinal nodes are involved as centers of incubation and dissemination in several 
diseases including lung cancer, tuberculosis, and anthrax.34-135-136 Treatment 
and control of these diseases is hard to accomplish because of the limited access of 
drugs to mediastinal nodes, using common pathways of drug delivery. Also, the 
anatomical location of mediastinal nodes represents a difficult target for external 
beam irradiation, due to its close proximity to major vessels and the heart. The use 
of the avidin/biotin-liposome method has been investigated as a carrier system 
for drug delivery to mediastinal nodes, using intrapleural injection as the pathway 
of delivery.137-138 Drug delivery, using the avidin/biotin-liposome system, injected 
intrapleurally could solve some of these limitations and offer several advantages 
for the treatment of these diseases. Only minimal investigation has been performed 
using intrapleural administration as a pathway for drug delivery to mediastinal 
nodes. Perez-Soler et ol. have investigated the intrapleural route of administration 
of liposome encapsulated drugs for the treatment of malignant pleural effusion in 
human patients.137 
Our group has performed studies to investigate the use of the avidin/biotinliposome 
system for targeting the mediastinal node following intrapleural administration 
in rats. Studies were performed by injecting biotin-liposomes into the 
pleural space following by an injection of avidin 2 hrs previously. By 22 hrs after 
injection, good retention (15.7% ID /mediastinal nodes; 515% ID/g) of liposomes 
was achieved in the mediastinal nodes with the avidin/biotin-liposome system. 
The scintigraphic image that visually demonstrates the mediastinal node uptake is 
shown in the image in the left panel of Fig. 4. An actual photograph of the blue stained 
mediastinal nodes obtained during necropsy is shown in Fig. 6. The images demonstrate 
the high uptake of liposomes in the mediastinal nodes. In the absence of avidin, 
liposomes were minimally retained in the nodes (< 1.0% ID/organ; 36% ID/g). This 
approach was the reverse of the sequence used in prior subcutaneous and intraperitoneal 
studies, in which avidin was injected after the biotin-liposomes.3-18 
The specific targeting of a liposome-encapsulated drug to mediastinal lymph 
nodes could result in a prolonged targeted sustained depot-like delivery of high 
drug concentrations to these nodes, while the liposomes are slowly degraded and 
metabolized by phagocytic cells located within these nodes. Future experiments 
using intrapleural injection of the avidin/biotin-liposome system to target drugs 
to mediastinal nodes should be pursued. 
586 Phillips 
Avidin Pleural 
Biotin-Liposomes-Pleural 
Avidin Pleural 
Biotin-Liposomes Peritoneal 
Fig. 4. Using the avidin-biotin liposome system in a rat model, high levels of liposomes 
were trapped in the mediastinal node when the avidin was injected in the pleural space 
and followed 2 hrs later by injection of the radiolabeled biotin-liposomes as demonstrated 
on the image on the left hand side. The avidin alone was injected in the pleural space and 
radiolabeled biotin liposomes were injected in the peritoneal space. Scintigraphic images 
were performed at 24 hrs. High levels of liposomes accumulated in the diaphragm as well 
as the mediastinal nodes. The diaphragm is the linear structure with intense uptake at the 
bottom of the chest region. 
12.5. Avidin biotin for diaphragm and mediastinal 
lymph node targeting 
Using the avidin-biotin liposome system, it was serendipitously discovered that 
when biotin-liposomes were injected into the peritoneal cavity and avidin was 
simultaneously injected into the pleural cavity, the liposomes aggregated strongly in 
the diaphragm as well as in the mediastinal nodes. This accumulation in the 
diaphragm occurred when the avidin draining from the pleural space into 
the diaphragmatic lymphatics encountered the biotin-liposomes draining from the 
peritoneal space, causing the liposomes to aggregate within the diaphragm. The 
scintigraphic image of this diaphragm and mediastinal node accumulation is shown 
in the image of the right panel in Fig. 4. The scintigraphic image shows the intense 
activity of linear uptake in the region of the diaphragm, as well as uptake in the 
mediastinal nodes. A photographic picture of the diaphragm is shown in Fig. 5. 
The blue dye containing biotin-liposomes accumulate in the linear lymphatic vessels 
coursing through the diaphragm. This study confirms the fact that in the rat, 
the pleural lymphatic drainage pathway and the peritoneal lymphatic drainage 
pathway share the same lymphatic vessels in the diaphragm. It is not known if 
there could be some useful applications for diaphragmatic drug delivery, but one 
potential application is for the treatment of mesothelioma. Methelioma is a cancer 
of the diaphragm that generally has a very poor prognosis.139 This prognosis has 
not changed with any attempted therapies including surgery, chemotherapy and 
radiation.139 
Nanoparticles for Targeting Lymphatics 587 
Fig. 5. A photographic picture of the diaphragm of a rat 24 hrs following injection of 
the biotin liposomes containing blue dye in the peritoneum and injection of avidin in the 
peritoneum corresponding to the right hand image on Fig. 4. The blue dye containing biotinliposomes 
accumulate in the linear lymphatic vessels coursing through the diaphragm. 
Fig. 6. A photographic picture of the biotin-liposome containing blue dye accumulating 
in the two mediastinal lymph nodes 24 hrs following injection of avidin in the pleural space 
followed by biotin-liposomes administration in the pleural space as corresponding to the 
left hand image of Fig. 4. The heart is between and forceps and the mediastinal nodes are 
just in front of the thymus. The intense blue staining of the mediastinal lymph nodes can 
clearly be seen at 24 hr post administration of the liposomes. 
13. Nanoparticles for Cancer Therapy 
13.1. Intralymphatic drug delivery to lymph nodes 
One of the first studies to investigate the possible use of drugs delivered intralymphatically 
was performed by Hirnle.1 This study investigated the anti cancer drug, 
588 Phillips 
Bleomycin, which was suspended in an oil suspension known as Oil Bleo. This Oil 
Bleo was injected directly into catheterized lymphatic vessels in dogs. The movement 
of this agent through the lymph nodes and lymphatic vessels was fairly rapid 
with peak drug concentrations reached in the blood 15 min after the intralymphatic 
administration of Oil Bleo. The drug entering the blood was considered to be a 
spillover from the lymphatic system. Spillover occurred because the drug moved 
completely through the lymphatic vessels and rejoined the circulation at the thoracic 
duct. Administering the drug this way required a very tedious catheterization 
process of the small lymphatic vessels of the extremities. Although drug concentrations 
were very high in the lymphatic vessels for a fairly short time, the retention 
of the oil emulsion in the lymphatics was minimal. 
Following this work with anticancer bleomycin oil emulsions, Hirnle turned to 
liposomes as an ideal carrier for intralymphatically delivered drugs.1 A study in rabbits 
used liposome-encapsulated Bleomycin in which the liposomes were injected 
directly into the lymphatic vessels of the hindlegs of rabbits. Lymph nodes were 
removed and measured for bleomycin content at various times following administration. 
Three days following intralymphatic administration, the drug concentration 
in the popliteal lymph nodes was 42 /zgram/gram of node. Drug deposition 
and apparent release was sustained over a very long period because concentrations 
of Bleomycin in the lymph nodes of 0.18 /xg/gram were measured in the popliteal 
nodes at one month following injection.1 
Further studies were performed by Hirnle with blue dye containing liposomes 
composed of 80% phosphatidylcholine and 20% cholesterol. The liposomes had 
a homogeneous size of approximately 170 nm in diameter. The total amount of 
blue dye injected was 1.6 mg. When the rabbits were sacrificed 28 days later, the 
retroperitoneal lymph nodes were visually blue and had a concentration of 172 /zg 
blue dye/gram of lymph node. Unfortunately, when these liposomes were administered 
by direct intralymphatic injection in the hindleg of a rabbit, a large fraction 
of the intact liposomes were found to spill over into the circulation. 
Several conclusions were derived from this endolymphatic research with liposomes 
directly infused into the lymphatic vessels. The amount of drug administered 
intralymphatically should not exceed that which would be administered 
intravenously, because of the large amount of liposomes that spill over from the 
lymphatics to the circulation. The limiting factor in administering drugs lymphatically 
is the amount of the therapeutic agent that moves completely through the 
lymphatic system and into the circulation through the thoracic duct. The tolerated 
amount of spillover should be considered in regard to the toxicity of these liposomal 
agents to the rest of the body. The volume used in humans should remain 
low, with no more than 4 ml of liposomes being administered into the canulated 
lymphatic vessels of each leg. It was also suggested that the drug will remain longer 
Nanoparticles for Targeting Lymphatics 589 
in the lymphatics, if the patient remains in bed for 1 day after endolymphatic liposome 
administration. Most importantly, the lymph nodes will still be filled with 
measurable amounts of drug a month after injection. Hirnle also introduced the 
concept that the prolonged retention of anticancer drugs in the lymphatics might 
be effective for the prevention of lymphatic metastasis.1 
13.2. Nanoparticles for tteatment of metastatic lymph nodes 
of upper Gl malignacies 
Much work has been performed by investigators in Japan to develop a novel drug 
delivery system for the treatment of lymph node metastasis from the cancers of the 
esophagus and stomach.140 This interest by Japanese researchers are likely to stem 
from the high incidence of upper gastrointestinal cancer in Japan. This work has 
also been stimulated by the fact that examinations of surgically resected specimens 
revealed that the cancer of the upper digestive tract metastasize to regional lymph 
nodes in 20-30% of patients, even when cancer invasion is limited to the mucosa or 
submucosa. This has led to the conclusion that even in patients with these superficial 
cancers, it is important to treat patients with gastric and esophageal cancer for 
potential lymph node metastasis. 
As an attempt to increase drug delivery to lymph nodes that drain from these 
upper digestive system cancers, a new activated carbon nanoparticle formulation of 
the anti cancer drug methotrexate was developed for the treatment of lymph node 
metastasis in patients with cancers of the upper digestive system.141 Methotrexate 
was mixed with 20 nm-sized activated carbon nanoparticles in a concentration of 
50 mg activated carbon/ml and polyvinylpyrrolidone. This made a suspension of 
methotrexate loaded carbon particles with an average size of 167 nm, as determined 
by photon correlation spectroscopy. Methotrexate was shown to be absorbed onto 
the activated carbon nanoparticles at a concentration of 25 mg/ml. 
Mouse leukemia P388 cells were used as an experimental tumor because 
subcutaneously implanted P388 was known to metastasize to lymph nodes within 
7 days.141 Experiments were carried out on day 7 when the cancer metastasis were 
known to be in the popliteal lymph nodes. In mice which received an inoculation 
of P388 leukemia cells and drug treatment using the same procedures, the 
treatment effects on metastases to the regional lymph nodes were significantly 
greater in mice treated with the methotrexate loaded activated carbon particles 
than in those given methotrexate aqueous solution. Methotrexate concentrations 
in the popliteal lymph nodes were 240 x 10~6 mol/kg, with the carbon particles 
at 1 hr versus 96 x 10~6 mol/kg with the free drug. This level rapidly dropped to 
18 x 10~6 mol/kg by 3 hrs versus 1.8 x 10~6 mol/kg for the free drug. By 6 hrs, levels 
of methotrexate were undetectable in the lymph node. Blood levels of drug in the 
590 Phillips 
serum also increased rapidly, indicating that the association of the drug with the 
carbon nanoparticles was not very stable. Blood levels at 30 min following injection 
were only slightly lower than those of subcutaneously injected free drug. These 
levels peaked at 30 min and were at very low concentration at 3 hrs and were non 
detectable in the serum at 6 hrs. This rapid drop in drug blood levels would suggest 
that the carbon particles were not as effective at binding the methotrexate, 
compared with other liposome-encapsulated drugs, reported to be measured in 
lymph nodes in detectable quantities for 1 month after subcutaneous injection.1 No 
studies were reported describing the stability of this drug attachment to the carbon 
nanoparticles during incubation in serum. Such studies are essential for the evaluation 
of the stability of the nanocarrier. Even with this small effect of improved 
delivery to lymph nodes, survival in this animal model was increased from 12 days 
in the non-treatment group to 17 days in the carbon particle methotrexate group, 
14 days in the free methotrexate group injected subcutaneously and 13 days in 
the free methotrexate group injected IV. This same group has gone on to perform 
clinical trials in human patients, in which methotrexate-carbon nanoparticles were 
injected locally for the treatment of cancer of the upper digestive tract.141 
This treatment approach has also been applied in small pilot studies by other 
groups in Japan, who have also reported survival benefit for the treatment of gastric 
cancer.142 Another group injected carbon nanoparticles with absorbed bleomycin 
for the treatment of esophageal cancer. In this study, bleomycin nanoparticles 
were injected into the esophageal cancer 3 days prior to surgery. Degenerative 
or inflammatory changes were microscopically observed in 6 of 23 lymph nodes, 
with metastatic foci indicating to these researchers that bleomycin carbon particles 
could be a useful tool in targeting chemotherapy for esophageal cancer.143 
This same activated carbon particle methodology has been applied to the 
treatment of other cancers. In one study, breast cancer patients were injected intratumorally 
and peritumorally, with aclarubicin absorbed to activated carbon nanoparticles 
or in free solution.144 Following this injection, the patient had surgery and 
the tumor and peritumoral tissue were removed as well as regional lymph nodes. 
Drug levels in the lymph nodes were shown to be significantly higher with the carbon 
nanoparticle associated drug, compared with that of free drug (42 /xg/g tissue 
versus 20 Mg/g tissue). 
In a recent study, local injection of mitomycin C bound to activated carbon 
(M-CH) combined with intraperitoneal hyperthermic hypo-osmolar infusion 
(IPHHOI) was intraoperatively administered to prevent lymph node recurrence 
and peritoneal recurrence of gastric cancer.145 The 1- and 2-year survival rates for 
the M-CH1 + IPHHOI group were 91.2 and 72.1%, and those for the control group 
were 78.9 and 45.5%. The M-CH1I + PHHOI group reaped a significant survival 
benefit (p = 0.0352) compared with the control group. Although this study was 
Nanoparticles for Targeting Lymphatics 591 
conducted in a small number of randomly selected patients with a short follow-up 
period, compared with the control group, the M-CH1 + IPHHOI group had a beneficial 
effect in preventing lymph node recurrence and peritoneal recurrence, after 
curative gastrectomy for advanced gastric cancer. 
13.3. Lessons from endolymphatic radioisotope therapy 
Much of the earlier work has been performed on the lymphatic delivery of the 
radiotherapeutic lipid emulsion, 1-131 (131I)-lipiodol, which is a lipid emulsion 
of iodinated ethylic ester of poppy-seed oil. 131I emits a therapeutic beta particle 
which is responsible for its therapeutic effects. This early work provides useful 
lessons for the potential delivery of radiotherapeutic nanoparticles into the lymphatics 
for the treatment of cancer. 131I endolymphatic therapy consists of direct 
infusion of 131I-lipiodol into the lymphatic vessels of the cancer affected extremity. 
Initially, endolymphatic isotope therapy had such promising early clinical results 
that the M.R.C. (Medical Research Council) in the U.K. set up a clinical trial in 
1966. This clinical trial compared patients with lower extremity melanoma who 
received 131I-lipiodol endolymphatic therapy to those who were treated with standard 
methods.146 Although there was no difference in the 5-year survival rate 
between the groups, lymph node recurrence was significantly different with only a 
2.3% lymph node recurrence rate with the 131I-lipiodol therapy, versus 19% lymph 
node recurrence rate with standard therapy. The conclusion from this study was 
that endolymphatic isotope therapy was justified in specialized centers where good 
results could be obtained.146 
Following this initial investigation, many other studies of endolymphatic radiotherapy 
were performed.147-149 Studies of radiation dosimetry found that the average 
radiation dose absorbed by the lymphatic tissues with this therapy was 90 rads. 
Unfortunately, this method was found to be limited by the hazard of radiation damage 
to the lungs.150 Approximately 80% of these patients had detectable concentrations 
of 131I radioactivity in the lung fields. The average radiation dose to the lungs 
was 299 rads. It is very evident that 131I-lipiodol becomes trapped in the lungs after 
re-entering the thoracic duct following therapy. This spillover from the lymphatic 
system that accumulates in the lungs, led to the recommendation that patients 
receiving this 131I-lipiodol endolymphatic therapy rest in bed for several days, so 
that the maximum amount of 131I would remain in the lymphatic vessels and not be 
pushed through to the lungs. It was this large lung radiation dose that eventually 
led to the discontinuation of these1311-lipiodol studies, even in the face of promising 
results for lymphatic therapy and the prevention of local lymphatic metastasis. 
Endolymphatic 131I-lipiodol therapy has been used to treat 426 patients 
with lymphoma. Traditional X-ray lymphography was performed during the 
592 Phillips 
administration of the therapeutic 131I-lipiodol. These studies found that endolymphatic 
therapy was not of value in cases where there was evidence of lymph nodes 
already involved with cancer at the time of the treatment. However, in cases where 
the lymphography was apparently negative, the 131I-lipiodol did produce a statistically 
significant reduced incidence of relapse in the inguino-retroperitoneal 
nodes.151 This suggests that the 131I-lipiodol therapy was effective in treating 
micrometastasis in the lymph nodes. 
One interesting study carried out with 131I-lipiodol examined the effect of prior 
external beam irradiation on lymph node uptake of endolymphatically infused 
iodinated 131I-lipiodol.152'153 In this study, 2 ml of 131I-lipiodol (76 mg of iodine per 
ml) was injected subcutaneously into 9 normal adult beagle dogs. Targeted lymph 
node groups were evaluated with computed tomography (CT). Lymph nodes were 
irradiated with 50 Gy in 25 fractions of 2 Gy per day, beginning 28-35 days after the 
CT examination. Contrast media administration and quantitative CT imaging were 
again performed 3 months after irradiation. Contrast material uptake resulted in a 
2-fold increase in node volume before irradiation (p < 0.0001). Prior to the external 
beam irradiation, mean attenuation of contrast-enhanced nodes increased to 230- 
330 Hounsfield units from a precontrast enhancement value of 36.5 Hounsfield 
units. After irradiation, opacified node volumes decreased to approximately 25%- 
50% of their preirradiation volumes (p < 0.02), but contrast material uptake in the 
lymph node only decreased by 10%-15% after irradiation. This uptake in the lymph 
node was not significantly less than the preirradiation uptake. Qualitatively, no 
substantial difference was found between irradiated and nonirradiated nodes. The 
external beam irradiation treatment decreased lymph node size, but the imaging 
characteristics of opacification were not otherwise appreciably altered 3 months 
after irradiation. A later study at 12 months showed slightly smaller lymph nodes 
and a lesser uptake of the subcutaneously injected 131I-lipiodol,152 however, the 
lymph nodes appeared to tolerate the 50 Gy dose without significant alteration in 
their function. 
14. Advantages of Nanoparticles for Lymphatic Radiotherapy 
Compared with the previously discussed1311-lipiodol emulsion, nanoparticles have 
many potential advantages as carriers of therapeutic radionuclides for endolymphatic 
therapy. These include the fact that nanoparticles do not accumulate in the 
lungs to any degree and the ability to control the release and the choice of the 
particular isotope that is attached or encapsulated in the nanoparticle. The high 
lung uptake of 131I-lipiodol is due to the lipophilic nature of its oil component 
causing it to be absorbed by the lungs, which is the first significant vascular capillary 
bed encountered after the 131I-lipiodol rejoins the circulation. It is well known 
that intravenously administered liposome nanoparticles or liposome nanoparticles, 
Nanoparticles for Targeting Lymphatics 593 
returning to the blood following drainage from the lymphatic system, do not accumulate 
in the lungs to any significant degree.18 
15. Intraoperative Radiotherapy for Positive Tumor Margins 
and for Treatment of Lymph Nodes 
One possible use of radiotherapeutic nanoparticles is to target residual tumor in 
the intraoperative situation. In many cases, the surgeon is unable to remove all of 
the cancer during surgery, so that the margins of the resected tumor are positive. 
This generally means that there is cancer remaining at the operative site which 
severely compromises patient's survival. This positive margin can frequently be 
determined during the operation. Radiotherapeutic nanoparticles that target residual 
tumor could be injected in the region of the positive tumor margin to sterilize the 
surgical margin of tumor cells. Since the radiotherapeutic nanoparticles will drain 
through the lymph nodes, they would also have the potential to treat micrometastasis 
in those nodes. Nanoparticles could therefore provide an additional tool for 
the surgeon, particularly when the margins of the tumor are positive. 
Even when the margins of the tumor are negative, there is frequently a reoccurrence 
of cancer in the local region or in the nodes that drain from the local region. 
Cancer surgeons spend many hours per surgery uncovering and removing lymph 
nodes in the region of the tumor carefully, without damaging other critical vessels 
and nerves. Although these surgeries are very long, it is not always possible to find 
and remove all of the lymph nodes in the local region of the tumor. Removal of distant 
lymph nodes that also receive lymph drainage from the tumor is usually not 
possible. The application of therapeutic nanoparticles intraoperatively could provide 
an additional tool to treat micrometastasis in lymph nodes, with the goal of 
decreasing local reoccurrences. Extensive clinical trials would have to be performed 
to determine the effectiveness of this approach, similar to those that have already 
been performed with 131I-lipiodol. Effective treatment of lymph nodes draining 
from a tumor could decrease the need for tedious surgical removal of lymph nodes. 
One possible method to ensure good lymph node targeting of nanoparticles in 
the intraoperative situation would be to use the avidin/biotin lymph node targeting 
system to ensure trapping of the particles in the lymph nodes that drain from 
the tumor. This methodology would also limit the spillover of radiotherapeutic 
nanoparticles from the lymphatic vessels into the bloodstream. 
16. Potential of Using Radiolabeled Nanoparticles for 
Intratumoral Radionuclide Therapy 
The direct injection of therapeutic agents into solid tumors has been recently 
investigated.154-157 These studies using direct injection of nanoparticles into tumor 
594 Phillips 
have used many different therapeutic agents. For instance, direct injection of 
nanoparticles into solid tumor have been investigated as a method of deliverying 
genes into tumors.158 This approach has also been applied in combination with 
external physical modalities. Magnetic nanoparticles have been directly injected 
into a solid tumor and exposed to alternating current as a new type of thermal 
ablation of solid tumors.157 
The particulate nature of nanoparticles appears to offer significant advantages 
for direct intratumoral administration. Nanoparticles appear to diffuse to some 
degree through the interstitial space of the tumor along primitive and chaotic lymph 
vessels within the tumor. The degree of diffusion may depend on the characteristics 
of the particular nanoparticle injected. Nanoparticle intratumoral diffusion should 
result in improved solid cancer therapy due to a more homogeneous distribution 
throughout the tumor. In spite of this potential for intratumoral diffusion, nanoparticles 
can still be well retained within the tumor. When free unencapsulated drug is 
injected intratumorally, it appears to be absorbed directly into the blood supply of 
the tumor with less diffusion through the tumor, so that there is a less homogeneous 
dose throughout the tumor following the intratumoral injection of a free drug, as 
compared with intratumoral injection of nanoparticles. In addition, depending on 
the nature of the free drug, free drug is likely to be cleared from the tumor rapidly 
by direct absorption into the tumor blood capillaries. 
Even with this improved local diffusion associated with nanoparticles compared 
with free drug, obtaining a homogeneous distribution throughout the solid 
tumor with intratumoral administration of nanoparticles still remains a challenge. 
One approach is to use modifications of the injection method such as multiple sites 
of injections within the solid tumor.154 This approach has recently been applied in 
the case of gene delivery with nanoparticles. Another possibility is the use of betaemitting 
therapeutic isotopes attached to nanoparticles. The beta-emissions penetrate 
millimeter distances away from the nanoparticle, enabling the beta-emitting 
nanoparticles to deliver therapy to regions of the solid tumor that the nanoparticles 
cannot reach themselves. Many other approaches to solve the problem of homogenous 
distribution within a solid tumor. Nanoparticles may be part of, but not likely 
the complete solution, to obtaining a very homogeneous distribution within a solid 
tumor. 
A significant advantage of nanoparticles for use in intratumoral injection is that 
they are more likely to move into the lymphatic vessels that drain from the solid 
tumor, where they have the chance to deliver anti cancer therapy to the sentinel 
lymph node and other lymphatics that drain from the tumor. Therefore, it is anticipated 
that this intratumoral injection would not only treat the tumor, but could 
also potentially treat lymph nodes that receive drainage from the tumor such as the 
sentinel node. These lymph nodes could possibly contain metastasis.159 
Nanoparticles for Targeting Lymphatics 595 
17. Liposome Pharmacokinetics after Intratumoral 
Administration 
Studies of liposome intratumoral pharmacokinetics have been stimulated by 
attempts to use liposomes as gene carriers. Clinical trials using cationic liposomes, 
carrying E1A gene, were performed to treat squamous cell carcinoma, using an 
intratumoral injection technique for intratumoral administration.160161 Pharmacokinetic 
studies have indicated that the size and surface charge of liposomes have 
a significant effect on their in vivo distribution.162,163 
Increasing the liposome diameter and adding a positive surface charge to the 
liposomes slowed their clearance from the injection site, compared with smallersized 
and neutral charged liposomes respectively. At 2 hrs after intratumoral injection, 
~ 70% and 90% of injected dose remained in the tumor with a 254.0 ± 5.1 nm 
neutral liposome and a 125.0 ± 29.4 nm cationic liposome respectively.163 Based on 
their observation of intratumorally administered cationic liposomes, Nomura et al. 
stated that there is a need to improve the control of the cationic liposome complexes 
to ensure a better distribution throughout the tumor.163 Biodistribution of 
111 In-labeled pegylated liposomes following intratumoral administration has also 
shown that liposomes have excellent potential as vehicles for intratumoral drug 
delivery.95 
18. Rhenium-Labeled Liposomes for Tumor Therapy 
Our group has developed a novel method of labeling liposomes with the 
radioisotope of rhenium. This method uses N,N-bis(2-mercaptoethyl)-N',N'- 
diethyl-ethylenediamine (BMEDA) to post-load either 99mTc, rhenium-188 (188Re) 
or rhenium-186 (186Re) into liposomes.159 
One of the significant advantages of rhenium-labeled nanoparticles that carry 
therapeutic beta particles is the short range field effect that they have, due to the 
fixed range of beta particle penetration (i.e. 2 mm for rhenium-186 and 4 mm for 
rhenium-188).164 This length of penetration is adequate to treat a large number of 
cancer cells in the region of the nanoparticle, but not so far as to cause extensive 
damage to normal tissue. The 2-4 mm range of beta emission penetration with 
the rhenium-186/188 isotopes compares favorably with 131I, which only has 1 mm 
average beta particle penetration combined with a high energy gamma photon. The 
4 mm treatment field with rhenium-188 is adequate for treating most lymph nodes, 
while limiting the dose to normal structures. This field effect of the beta particle can 
compensate, to some degree, for a heterogeneous distribution of the nanoparticles 
within cancer containing lymph nodes. The nanoparticle simply has to reach within 
a 4 mm vicinity of the cancer cells.164 
596 Phillips 
For every 10 beta emissions, both rhenium isotopes, rhenium-186 and rhenium- 
188, emit a single gamma photon. This is an ideal ratio of beta to gamma emissions. 
A higher number of gamma emissions would deliver an excessive dose outside the 
local region of the tumor, as is the case for1311 which has a 1:1 ratio of beta particles 
to gamma photons. The photon emission energy of both rhenium isotopes is in the 
range of the photon energy of 99mTc (140 keV), so that the radiolabeled nanoparticles 
can be tracked through the body as they move through the lymphatic vessels. 
Many therapeutic radioisotopes are pure beta emitter, so that it is more difficult to 
track their distribution in the body. Rhenium has also many other advantages over 
most heavy metal radiotherapeutic isotopes, such as yittrium-91, because it has 
almost no affinity for bone uptake. It shares this characteristic with 99mTc, as both 
radioisotopes tend to be cleared through the kidney, while most heavy metal betaemitting 
radioisotopes have a high affinity for bone. This high bone accumulation 
can deliver a high radiation dose to bone marrow cells which are very sensitive to 
radiation. This occurs when the radioisotope becomes separated from its chelator, 
following metabolism in the body. 
Previous theoretical dosimetry studies have addressed the potential use 
of radiotherapeutic liposomes for the treatment of tumors via intravenous 
injection.165-167 In addition to these intravenous investigations, our group has 
investigated the potential use of rhenium-liposomes for intratumoral therapy.159 
There are some significant advantages of using the intratumoral delivery route for 
rhenium-liposomes compared with intravenous injection, such as the much lower 
radiation dose delivered to liver, spleen, kidney and other normal tissues, and the 
potential of simultaneous targeting of metastatic lymph nodes that drain from the 
region of the tumor.3 
99mTc-liposomes can be used to pre-evaluate the suitability of using 186Re/ 188Reliposomes 
to treat a tumor. This is because the same chemistry is used to label liposomes 
with the diagnostic isotope, 99mTc, as the therapeutic rhenium isotopes. The 
likely dose distribution from the rhenium-liposomes can be calculated by performing 
SPECT/CT images of the 99mTc-liposome distribution, in order to determine the 
potential dose distribution of the rhenium-liposomes.91 
We have performed studies with 99mTc to assess intratumoral administration of 
radiolabeled liposomes. In these studies, prolonged tumor retention and very high 
tumor-to-normal tissue ratio of 99mTc-activity were observed (manuscript submitted 
for publication). 99mTc-liposomes were injected intratumorally into a head and neck 
tumor in a rat model, using the same methodology for labeling liposomes with 
radiotherapeutic rhenium. 99rnTc-liposomes had good tumor retention with 47.6 to 
65.7% of injected activity still remaining in the solid tumors at 44 hrs after injection, 
while unencapsulated 99mTc-BMEDA cleared from tumors quickly, with only 37.1 % 
of injected activity remaining in tumors at 2 hrs and 19.4% at 44 hrs. 
Nanoparticles for Targeting Lymphatics 597 
19. Nanoparticles for Immune Modulation 
A few very preliminary studies suggest that the delivery of therapeutic betaemitting 
radioisotopes to lymph nodes has the potential to modulate the immune 
system for therapeutic benefit of auto-immune disease and for the induction of 
tolerance in organ transplantation. These preliminary studies suggest the possibility 
that beta-emitters delivered to lymph nodes results in a decreased immune 
response in the organs and regions of the body that drain that lymph node. This 
decreased immune response has been demonstrated in pilot studies of patients 
with rheumatoid arthritis, as well as in patients that have received transplanted 
kidneys. 
In one study, a method was developed and tested for the treatment of patients 
with rheumatoid arthritis, using radiotherapeutic beta-emitting gold-198 colloid 
particles which were infused into the lower limb lymphatic vessels. More than 
50 patients were treated. A positive therapeutic effect was observed in 84% of 
the treated patients. This endolymphatic radiotherapy with gold colloid particles 
made it possible to give up cytostatic and glucocorticoid medications and to 
reduce the dosage of nonsteroid anti-inflammatory drugs.168 Immune modulation 
by radioparticle accumulation in the lymph nodes could also explain some of the 
beneficial effects of radiation synovectomy. In this procedure, radiolabeled particles 
that emit beta particles are injected into the joints of patients with rheumatoid 
arthritis.169 
A second study also provides evidence of tolerance induction by the pre- transplant 
endolymphatic infusion of 131I-lipiodol. This procedure was performed as 
a pre-transplant preparation for patients receiving a kidney transplant. Twenty 
six years later, the outcome in the patients that received the 131I-lipiodol was 
compared with that of another group of patients that did not receive the 131Ilipiodol 
therapy, but were treated with a standard maintenance dose of azathioprine. 
The incidence of rejection crises was greatly reduced in the group that 
received the 131I-lipiodol therapy, compared with the standard treatment group 
(21% versus 74%, p = 0.003). The authors of this study concluded that the pretransplant 
treatment with 131I-lipiodol had an extended immunosuppressive effect 
and could be indicated for cadaveric renal allograft recipients, especially those 
showing high panel reactivity. It was also relatively innocuous, as there was no 
compromising of either the thyroid gland or the gonad function and there was 
no increase in tumor incidence in these patients over the 26-year period.170,171 
Local infusion of nanoparticles carrying therapeutic beta-emitting radioisotopes 
that targeted the lymph nodes might have potential applications for the prevention 
of transplanted organ rejection, as well as the treatment of auto-immune 
disorders. 
598 Phillips 
20. Conclusions 
The delivery of nanoparticles to lymph nodes for therapeutic purposes is promising. 
Significant progress has been made in understanding the various processes 
involved in nanoparticle delivery and in the development of potential systems for 
targeting nanoparticles to lymph nodes. Lymph node delivery appears promising 
for improving cancer and infectious disease therapy, treatment of autoimmune disease 
and for improvement of vaccine systems. 
Acknowledgments 
The author is grateful to Dr. Beth Goins for her help and critical reading of the 
manuscript. 
References 
1. Hirnle P (1997) Liposomes for drug targeting in the lymphatic system. Hybridoma 
16:127-32. 
2. Wisner ER et al. (2003) Sentinel node detection using contrast-enhanced power Doppler 
ultrasound lymphography. Invest Radiol 38:358-65. 
3. Phillips WT, Klipper R and Goins B (2000) Novel method of greatly enhanced delivery 
of liposomes to lymph nodes. / Pharmacol Exp Ther 295:309-13. 
4. Moghimi SM (2003) Modulation of lymphatic distribution of subcutaneously injected 
poloxamer 407-coated nanospheres: The effect of the ethylene oxide chain configuration. 
FEBS Lett 540:241^. 
5. Oussoren C and Storm G (2001) Liposomes to target the lymphatics by subcutaneous 
administration. Adv Drug Del Rev 50:143-56. 
6. Duzgunes N et al. (2001) Enhanced inhibition of HIV-1 replication in macrophages 
by antisense oligonucleotides, ribozymes and acyclic nucleoside phosphonate analogs 
delivered in pH-sensitive liposomes. Nucleos Nucleot Nucleic Acids 20:515-23. 
7. Swartz MA (2001) The physiology of the lymphatic system. Adv Drug Del Rev 50:3-20. 
8. Porter CJ (1997) Drug delivery to the lymphatic system. Crit Rev Ther Drug Can Syst 
14:333-93. 
9. von Andrian UH and Mempel TR (2003) Homing and cellular traffic in lymph nodes. 
Nat Rev Immunol 3:867-78. 
10. Swartz MA and Skobe M (2001) Lymphatic function, lymphangiogenesis, and cancer 
metastasis. Microsc Res Tech 55:92-9. 
11. Hawley A, Davis S and Ilium L (1995) Targeting of colloids to lymph nodes: Influence 
of lymphatic physiology and colloidal charateristics. Adv Drug Del Rev 17:129-148. 
12. Tran L et al. (1993) Lymphatic drainage of hypertonic solution from peritoneal cavity 
of anesthetized and conscious sheep. / Appl Physiol 74:859-67. 
13. Papisov M and Weissleder R (1996) Drug delivery to lymphatic tissue. Crit Rev Ther 
Drug Can Syst 13:57-84. 
Nanoparticles for Targeting Lymphatics 599 
14. Ioachim HL (1982) Lymph Node Biopsy. JB Lippincott, Philadelphia, PA. 
15. Dixon TC, Meselson M, Guillemin J and Hanna PC (1999) Anthrax. N Engl } Med 
341:815-26. 
16. Phillips WT et al. (2001) Evaluation of [(99m)Tc] liposomes as lymphoscintigraphic 
agents: Comparison with [(99m)Tc] sulfur colloid and [(99m)Tc] human serum albumin. 
Nucl Med Biol 28:435-44. 
17. Moghimi SM and Rajabi-Siahboomi R (1996) Advanced colloid-based systems for efficient 
delivery of drugs and diagnostic agents to the lymphatic tissues. Prog Biophys Mol 
Biol 65:221^9. 
18. Phillips WT, Medina LA, Klipper R and Goins B (2002) A novel approach for the 
increased delivery of pharmaceutical agents to peritoneum and associated lymph 
nodes. / Pharmacol Exp Ther 303:11-6. 
19. Porter CJ and Charman WN (2001) Transport and absorption of drugs via the lymphatic 
system. Adv Drug Del Rev 50:1-2. 
20. Hanahan D and Weinberg RA (2000) The hallmarks of cancer. Cell 100:57-70. 
21. Chen C et al. (2004) Outcome after treatment of patients with mammographically 
occult, magnetic resonance imaging-detected breast cancer presenting with axillary 
lymphadenopathy. Clin Breast Cancer 5:72-7. 
22. Torabi M, Aquino SL and Harisinghani MG (2004) Current concepts in lymph node 
imaging. / Nucl Med 45:1509-18. 
23. Busby JE and Evans CP (2004) Old friends, new ways: Revisiting extended lymphadenectomy 
and neoadjuvant chemotherapy to improve outcomes. Curr Opin Urol 
14:251-7. 
24. Gagne JF, Desormeaux A, Perron S, Tremblay MJ and Bergeron MG (2002) Targeted 
delivery of indinavir to HIV-1 primary reservoirs with immunoliposomes. Biochim Biophys 
Acta 1558:198-210. 
25. Cohen OJ, Pantaleo G, Lam GK and Fauci AS (1997) Studies on lymphoid tissue from 
HIV-infected individuals: Implications for the design of therapeutic strategies. Springer 
Semin Immunopathol 18:305-22. 
26. Mand S, Debrah A, Batsa L, Adjei O and Hoerauf A (2004) Reliable and frequent detection 
of adult Wuchereria bancrofti in Ghanaian women by ultrasonography. Trop Med 
Lnt Health 9:1111-4. 
27. Reddy GS, Das LK and Pani SP (2004) The preferential site of adult Wuchereria bancrofti: 
An ultrasound study of male asymptomatic microfilaria carriers in Pondicherry, India. 
Natl Med } India 17:195-6. 
28. Dreyer G, Figueredo-Silva J, Carvalho K, Amaral F and Ottesen EA (2001) Lymphatic 
filariasis in children: Adenopathy and its evolution in two young girls. Am J Trop Med 
Hyg 65:204-7. 
29. Soulard R, Guigay J, Legal De Kerangal X and Saint-Blancard P (2001) [Mediastinal 
lymphatic filariasis]. Ann Pathol 21:431-4. 
30. El Setouhy M et al. (2004) A randomized clinical trial comparing single- and multidose 
combination therapy with diethylcarbamazine and albendazole for treatment of 
bancroftian filariasis. Am J Trop Med Hyg 70:191-6. 
600 Phillips 
31. Guarner J et al. (2003) Pathology and pathogenesis of bioterrorism-related inhalational 
anthrax. Am f Pathol 163:701-9. 
32. Jernigan JA et al. (2001) Bioterrorism-related inhalational anthrax: The first 10 cases 
reported in the United States. Emerg Infect Dis 7:933^4. 
33. Barakat LA etal. (2002) Fatal inhalational anthrax in a 94-year-old Connecticut woman. 
Jama 287:863-8. 
34. Ilgazli A, Boyaci H, Basyigit I and Yildiz F (2004) Extrapulmonary tuberculosis: Clinical 
and epidemiologic spectrum of 636 cases. Arch Med Res 35:435^1. 
35. Mukherjee JS et al. (2004) Programmes and principles in treatment of multidrugresistant 
tuberculosis. Lancet 363:474-81. 
36. Bermudez LE (1994) Use of liposome preparation to treat mycobacterial infections. 
Immunobiology 191:578-83. 
37. Zinkernagel RM et al. (1997) Antigen localisation regulates immune responses in a doseand 
time-dependent fashion: A geographical view of immune reactivity. Immunol Rev 
156:199-209. 
38. Schijns VE (2001) Induction and direction of immune responses by vaccine adjuvants. 
Crit Rev Immunol 21:75-85. 
39. Johansen P et al. (2005) Direct intralymphatic injection of peptide vaccines enhances 
immunogenicity. Eur J Immunol 35:568-574. 
40. Gregoriadis G, Bacon A, Caparros-Wanderley W and McCormack B (2002) A role for 
liposomes in genetic vaccination. Vaccine 20(Suppl 5):Bl-9. 
41. Maloy KJ et al. (2001) Intralymphatic immunization enhances DNA vaccination. Proc 
Natl Acad Sci USA 98:3299-303. 
42. Fokin AA, Robicsek F, Masters TN, Schmid-Schonbein GW and Jenkins SH (2000) 
Propagation of viral-size particles in lymph and blood after subcutaneous inoculation. 
Microcirculation 7:193-200. 
43. Eyles JE, Bramwell VW, Williamson ED and Alpar HO (2001) Microsphere translocation 
and immunopotentiation in systemic tissues following intranasal administration. 
Vaccine 19:4732-42. 
44. Phillips WT, Klipper R and Goins B (2001) Use of (99m)Tc-labeled liposomes encapsulating 
blue dye for identification of the sentinel lymph node. / Nucl Med 42:446-51. 
45. Oussoren C, Zuidema J, Crommelin DJ and Storm G (1997) Lymphatic uptake and 
biodistribution of liposomes after subcutaneous injection. II. Influence of liposomal 
size, lipid compostion and lipid dose. Biochim Biophys Acta 1328:261-72. 
46. Moghimi SM et al. (1994) Surface engineered nanospheres with enhanced drainage 
into lymphatics and uptake by macrophages of the regional lymph nodes. FEBS Lett 
344:25-30. 
47. Oussoren C and Storm G (1997) Lymphatic uptake and biodistribution of liposomes 
after subcutaneous injection: III. Influence of surface modification with poly(ethyleneglycol). 
Pharm Res 14:1479-84. 
48. Trubetskoy VS, Whiteman KR, Torchilin VP and Wolf GL (1998) Massage-induced 
release of subcutaneously injected liposome-encapsulated drugs to the blood. / Control 
Rel 50:13-9. 
Nanoparticles for Targeting Lymphatics 601 
49. Kuypers FA and de Jong K (2004) The role of phosphatidylserine in recognition and 
removal of erythrocytes. Cell Mol Biol (Noisy-le-gmnd) 50:147-58. 
50. Van Rooijen N, Kors N, vd Ende M and Dijkstra CD (1990) Depletion and repopulation 
of macrophages in spleen and liver of rat after intravenous treatment with liposomeencapsulated 
dichloromethylene diphosphonate. Cell Tissue Res 260:215-22. 
51. Van Rooijen N and Sanders A (1994) Liposome mediated depletion of macrophages: 
Mechanism of action, preparation of liposomes and applications. / Immunol Meth 
174:83-93. 
52. Oussoren C and Storm G (1999) Role of macrophages in the localisation of liposomes 
in lymph nodes after subcutaneous administration. Int J Pharm 183:37-41. 
53. Velinova M, Read N, Kirby C and Gregoriadis G (1996) Morphological observations 
on the fate of liposomes in the regional lymph nodes after footpad injection into rats. 
Biochim Biophys Acta 1299:207-15. 
54. Wisner ER et al. (1995) Indirect computed tomography lymphography of subdiaphragmatic 
lymph nodes using iodinated nanoparticles in normal dogs. Acad Radiol 
2:405-12. 
55. Wisner ER et al. (1994) Iodinated nanoparticles for indirect computed tomography 
lymphography of the craniocervical and thoracic lymph nodes in normal dogs. Acad 
Radiol 1:377-84. 
56. Wisner ER et al. (1996) Indirect computed tomography lymphography using iodinated 
nanoparticles to detect cancerous lymph nodes in a cutaneous melanoma model. Acad 
Radiol 3:40-8. 
57. Wolf GL et al. (1994) Percutaneous computed tomographic lymphography of normal, 
inflamed, and cancerous nodes in the rabbit. Invest Radiol 29(Suppl 2):S30-2. 
58. Mclntire GL et al. (2000) Time course of nodal enhancement with CT X-ray nanoparticle 
contrast agents: Effect of particle size and chemical structure. Invest Radiol 
35:91-6. 
59. Ketai LH et al. (1999) CT imaging of intrathoracic lymph nodes in dogs with bronchoscopically 
administered iodinated nanoparticles. Acad Radiol 6:49-54. 
60. Herborn CU et al. (2002) Interstitial MR lymphography with MS-325: Characterization 
of normal and tumor-invaded lymph nodes in a rabbit model. A]R Am } Roentgenol 
179:1567-72. 
61. Muldoon LL et al. (2004) Trafficking of superparamagnetic iron oxide particles (Combidex) 
from brain to lymph nodes in the rat. Neuropathol Appl Neurobiol 30:70-9. 
62. Nguyen BC et al. (1999) Multicenter clinical trial of ultrasmall superparamagnetic iron 
oxide in the evaluation of mediastinal lymph nodes in patients with primary lung 
carcinoma. / Magn Reson Imaging 10:468-73. 
63. Moghimi SM and Bonnemain B (1999) Subcutaneous and intravenous delivery of diagnostic 
agents to the lymphatic system: Applications in lymphoscintigraphy and indirect 
lymphography. Adv Drug Del Rev 37:295-312. 
64. Szebeni J (2001) Complement activation-related pseudoallergy caused by liposomes, 
micellar carriers of intravenous drugs, and radiocontrast agents. Crit Rev Ther Drug 
Can Syst 18:567-606. 
602 Phillips 
65. Bordat Cetal. (2000) Distribution of iron oxide nanoparticles in rat lymph nodes studied 
using electron energy loss spectroscopy (EELS) and electron spectroscopic imaging 
(ESI). JMagn Reson Imaging 12:505-9. 
66. Bellin MF et al. (1998) Lymph node metastases: Safety and effectiveness of MR imaging 
with ultrasmall superparamagnetic iron oxide particles — initial clinical experience. 
Radiology 207:799-808. 
67. Koh DM et al. (2004) Rectal cancer: Mesorectal lymph nodes at MR imaging with USPIO 
versus histopathologic findings — initial observations. Radiology 231:91-9. 
68. Harisinghani MG and Weissleder R (2004) Sensitive, Noninvasive Detection of Lymph 
Node Metastases. PLoS Med 1, e66. 
69. Morton DL et al. (1992) Technical details of intraoperative lymphatic mapping for early 
stage melanoma. Arch Surg 127:392-9. 
70. Focht SL (1999) Lymphatic mapping and sentinel lymph node biopsy. Aorn } 
69:802-9. 
71. Gradilone A et al. (2004) Detection of melanoma cells in sentinel lymph nodes by 
reverse transcriptase-polymerase chain reaction: Prognostic significance. Ann Surg 
Oncol 11:983-7. 
72. Wong JH, Cagle LA and Morton DL (1991) Lymphatic drainage of skin to a sentinel 
lymph node in a feline model. Ann Surg 214:637-41. 
73. Alex JC and Krag DN (1993) Gamma-probe guided localization of lymph nodes. Surg 
Oncol 2:137-43. 
74. Cody HS, 3rd et al. (2001) Complementarity of blue dye and isotope in sentinel node 
localization for breast cancer: Univariate and multivariate analysis of 966 procedures. 
Ann Surg Oncol 8:13-9. 
75. Hung JC et al. (1995) Filtered technetium-99m-sulfur colloid evaluated for lymphoscintigraphy. 
/ Nucl Med 36:1895-901. 
76. Martin RC, 2nd et al. (2000) Practical guidelines for optimal gamma probe detection of 
sentinel lymph nodes in breast cancer: Results of a multi-institutional study. For the 
University of Louisville Breast Cancer Study Group. Surgery 128:139^4. 
77. KapteijnBAef al. (1996) Reproducibility of lymphoscintigraphy for lymphatic mapping 
in cutaneous melanoma. / Nucl Med 37:972-5. 
78. Temple CL et al. (2000) Sentinel node biopsy in melanoma using technetium-99m rhenium 
colloid: The London experience. Ann Plast Surg 45:491-9. 
79. Hodgson Netal. (2001) A new radiocolloid for sentinel node detection in breast cancer. 
Ann Surg Oncol 8:133-7. 
80. Hirnle P, Harzmann R and Wright JK (1988) Patent blue V encapsulation in liposomes: 
Potential applicability to endolympatic therapy and preoperative chromolymphography. 
Lymphology 21:187-9. 
81. Dieter M, Schubert R and Hirnle P (2003) Blue liposomes for identification of the sentinel 
lymph nodes in pigs. Lymphology 36:39^17. 
82. Pump B and Hirnle P (1996) Preoperative lymph-node staining with liposomes containing 
patent blue violet. A clinical case report. / Pharm Pharmacol 48:699-701. 
Nanoparticles for Targeting Lymphatics 603 
83. Plut EM, Hinkle GH, Guo W and Lee RJ (2002) Kit formulation for the preparation 
of radioactive blue liposomes for sentinel node lymphoscintigraphy. / Pharm Sci 
91:1717-32. 
84. Michalet X et al. (2005) Quantum dots for live cells, in vivo imaging, and diagnostics. 
Science 307:538-44. 
85. Soltesz EG et al. (2005) Intraoperative sentinel lymph node mapping of the lung using 
near-infrared fluorescent quantum dots. Ann Thorac Surg 79:269-77; discussion 269-77. 
86. Liu J, Raveendran P, Shervani Z, Ikushima Y and Hakuta Y (2005) Synthesis of Ag and 
Agl Quantum Dots in AOT-Stabilized Water-in-CO(2) Microemulsions. Chemistry. 
87. Postema M, van Wamel A, Lancee CT and de Jong N (2004) Ultrasound-induced encapsulated 
microbubble phenomena. Ultrasound Med Biol 30:827-40. 
88. Dijkmans PA et al. (2004) Microbubbles and ultrasound: From diagnosis to therapy. Eur 
} Echocardiogr 5:245-56. 
89. Stride E and Saffari N (2004) The potential for thermal damage posed by microbubble 
ultrasound contrast agents. Ultrasonics 42:907-13. 
90. Hasegawa BH et al. (2002) Dual-modality imaging of cancer with SPECT/CT. Technol 
Cancer Res Treat 1:449-58. 
91. Wagner A et al. (2004) SPECT-CT for topographic mapping of sentinel lymph nodes 
prior to gamma probe-guided biopsy in head and neck squamous cell carcinoma. 
/ Craniomaxillofac Surg 32:343-9. 
92. Lopez R et al. (2004) Multimodal image registration for localization of sentinel nodes 
in head and neck squamous cell carcinoma. / Oral Maxillofac Surg 62:1497-504. 
93. Kumar R, Mavi A, Bural G and Alavi A (2005) Fluorodeoxyglucose-PET in the management 
of malignant melanoma. Radiol Clin North Am 43:23-33. 
94. Iyengar S, Chin B, Margolick JB, Sabundayo BP and Schwartz DH (2003) Anatomical 
loci of HIV-associated immune activation and association with viraemia. Lancet 
362:945-50. 
95. Harrington KJ et al. (2000) Pegylated liposomes have potential as vehicles for intratumoral 
and subcutaneous drug delivery. Clin Cancer Res 6:2528-37. 
96. Hawley AE, Ilium L and Davis SS (1997) Lymph node localisation of biodegradable 
nanospheres surface modified with poloxamer and poloxamine block co-polymers. 
FEBS Lett 400:319-23. 
97. Peyre M, Fleck R, Hockley D, Gander B and Sesardic D (2004) In vivo uptake of an experimental 
microencapsulated diphtheria vaccine following sub-cutaneous immunisation. 
Vaccine 22:2430-7. 
98. Trubetskoy VS, Frank-Kamenetsky MD, Whiteman KR, WolfGL and Torchilin VP (1996) 
Stable polymeric micelles: Lymphangiographic contrast media for gamma scintigraphy 
and magnetic resonance imaging. Acad Radiol 3:232-8. 
99. Torchilin VP (2004) Targeted polymeric micelles for delivery of poorly soluble drugs. 
Cell Mol Life Sci 61:2549-59. 
100. Segal AW, Gregoriadis G and Black CD (1975) Liposomes as vehicles for the local release 
of drugs. Clin Sci Mol Med 49:99-106. 
604 Phillips 
101. Osborne MP, Richardson VJ, Jeyasingh K and Ryman BE (1979) Radionuclide-labelled 
liposomes — a new lymph node imaging agent. Int J Nucl Med Biol 6:75-83. 
102. Patel HM, Boodle KM and Vaughan-Jones R (1984) Assessment of the potential uses 
of liposomes for lymphoscintigraphy and lymphatic drug delivery. Failure of 99mtechnetium 
marker to represent intact liposomes in lymph nodes. Biochim Biophys Acta 
801:76-86. 
103. Love WG, Amos N, Williams BD and Kellaway IW (1989) Effect of liposome surface 
charge on the stability of technetium (99mTc) radiolabelled liposomes. / Microencapsul 
6:105-13. 
104. Andersen AR, Friberg H, Lassen NA, Kristensen K and Neirinckx RD (1987) Serial 
studies of cerebral blood flow using 99Tcm-HMPAO: A comparison with 133Xe. Nucl 
Med Commun 8:549-57. 
105. Phillips WT et al. (1992) A simple method for producing a technetium-99m-labeled 
liposome which is stable in vivo. Int J Rad Appl Instrum B 19:539-47. 
106. Goins BA and Phillips WT in Liposomes: A Practical Approach, Torchilin VP and 
Weissig V (eds.) Oxford University Press, Oxford, 2003, pp. 319-336. 
107. Oussoren C et al. (1998) Lymphatic uptake and biodistribution of liposomes after subcutaneous 
injection. IV. Fate of liposomes in regional lymph nodes. Biochim Biophys Acta 
1370:259-72. 
108. Mangat S and Patel HM (1985) Lymph node localization of non-specific antibody-coated 
liposomes. Life Sci 36:1917-25. 
109. Wu MS, Robbins JC, Bugianesi RL, Ponpipom MM and Shen TY (1981) Modified 
in vivo behavior of liposomes containing synthetic glycolipids. Biochim Biophys Acta 
674:19-29. 
110. Medina LA, Calixto SM, Klipper R, Phillips WT and Goins B (2004) Avidin/biotinliposome 
system injected in the pleural space for drug delivery to mediastinal lymph 
nodes. / Pharm Sci 93:2595-608. 
111. Medina LA, Klipper R, Phillips WT and Goins B (2004) Pharmacokinetics and biodistribution 
of [llllnj-avidin and [99mTc]-biotin-liposomes injected in the pleural space 
for the targeting of mediastinal nodes. Nucl Med Biol 31:41-51. 
112. Fielding RM, Moon-McDermott L and Lewis RO (1999) Bioavailability of a small 
unilamellar low-clearance liposomal amikacin formulation after extravascular 
administration. / Drug Targ 6:415-26. 
113. Duzgunes N et al. (1988) Enhanced effect of liposome-encapsulated amikacin on 
Mycobacterium avium-M. intracellulare complex infection in beige mice. Antimicrob 
Agents Chemother 32:1404-11. 
114. Kinman L et al. (2003) Lipid-drug association enhanced HIV-1 protease inhibitor indinavir 
localization in lymphoid tissues and viral load reduction: A proof of concept 
study in HIV-2287-infected macaques. / Acquir Immune Defic Syndr 34:387-97. 
115. Dufresne I et al. (1999) Targeting lymph nodes with liposomes bearing anti-HLA-DR 
Fab' fragments. Biochim Biophys Acta 1421:284-94. 
116. Alberts DS et al. (2002) Intraperitoneal therapy for stage III ovarian cancer: A therapy 
whose time has come! / Clin Oncol 20:3944-6. 
Nanoparticles for Targeting Lymphatics 605 
117. Markman M (2003) Intraperitoneal antineoplastic drug delivery: Rationale and results. 
Lancet Oncol 4:277-83. 
118. Markman M (2004) Intraperitoneal hyperthermic chemotherapy as treatment of 
peritoneal carcinomatosis of colorectal cancer. / Clin Oncol 22:1527; author reply 
1529. 
119. Conti M, De Giorgi U, Tazzari V, Bezzi F and Baccini C (2004) Clinical pharmacology 
of intraperitoneal cisplatin-based chemotherapy. / Chemother 16(Suppl 5):23-5. 
120. Parker RJ, Hartman KD and Sieber SM (1981) Lymphatic absorption and tissue disposition 
of liposome-entrapped [14C]adriamycin following intraperitoneal administration 
to rats. Cancer Res 41:1311-7. 
121. Parker RJ, Priester ER and Sieber SM (1982) Comparison of lymphatic uptake, metabolism, 
excretion, and biodistribution of free and liposome-entrapped [14C]cytosine 
beta-D-arabinofuranoside following intraperitoneal administration to rats. Drug Metab 
Dispos 10:40-6. 
122. Rosa P and Clementi F (1983) Absorption and tissue distribution of doxorubicin 
entrapped in liposomes following intravenous or intraperitoneal administration. Pharmacology 
26:221-9. 
123. Morice P et al. (2004) Are nodal metastases in ovarian cancer chemoresistant lesions? 
Analysis of nodal involvement in 105 patients treated with preoperative chemotherapy. 
Eur J Gynaecol Oncol 25:169-74. 
124. Montero CA, Gimferrer JM, Baldo X and Ramirez J (2000) Mediastinal metastasis of 
ovarian carcinoma. Eur J Obstet Gynecol Reprod Biol 91:199-200. 
125. Zakaria ER, Simonsen O, Rippe A and Rippe B (1996) Transport of tracer albumin from 
peritoneum to plasma: Role of diaphragmatic, visceral, and parietal lymphatics. Am } 
Physiol 270:H1549-56. 
126. Yuan ZY, Rodela H, Hay JB, Oreopoulos D and Johnston MG (1994) 51Cr-RBCs and 
1251-albumin as markers to estimate lymph drainage of the peritoneal cavity in sheep. 
JAppl Physiol 76:867-74. 
127. Flessner MF, Parker RJ and Sieber SM (1983) Peritoneal lymphatic uptake of fibrinogen 
and erythrocytes in the rat. Am J Physiol 244:H89-96. 
128. Hirano K and Hunt CA (1985) Lymphatic transport of liposome-encapsulated 
agents: Effects of liposome size following intraperitoneal administration. / Pharm Sci 
74:915-21. 
129. Sharma A, Sharma US and Straubinger RM (1996) Paclitaxel-liposomes for intracavitary 
therapy of intraperitoneal P388 leukemia. Cancer Lett 107:265-72. 
130. Markman M et al. (1992) Phase I trial of intraperitoneal taxol: A Gynecoloic Oncology 
Group study. / Clin Oncol 10:1485-91. 
131. Lee MJ et al. (2002) Intraperitoneal gene delivery mediated by a novel cationic liposome 
in a peritoneal disseminated ovarian cancer model. Gene Ther 9:859-66. 
132. Hribaschek A et al. (2001) Prophylaxis of peritoneal carcinomatosis in experimental 
investigations. Int J Colorectal Dis 16:340-5. 
133. Hribaschek A et al. (2003) Intraperitoneal treatment using taxol is effective for experimental 
peritoneal carcinomatosis in a rat model. Oncol Rep 10:1793-8. 
606 Phillips 
134. Eltabbakh GH, Piver MS, Hempling RE, Recio FO and Intengen ME (1999) Clinical 
picture, response to therapy, and survival of women with diffuse malignant peritoneal 
mesothelioma. / Surg Oncol 70:6-12. 
135. Grinberg LM, Abramova FA, Yampolskaya OV, Walker DH and Smith JH (2001) Quantitative 
pathology of inhalational anthrax I: Quantitative microscopic findings. Mod 
Pathol 14:482-95. 
136. Jackson PJ et al. (1998) PCR analysis of tissue samples from the 1979 Sverdlovsk anthrax 
victims: The presence of multiple Bacillus anthracis strains in different victims. Proc Natl 
Acad Set USA 95:1224-9. 
137. Perez-Soler R et al. (1997) Phase I clinical and pharmacological study of liposomeentrapped 
NDDP administered intrapleurally in patients with malignant pleural effusions. 
Clin Cancer Res 3:373-9. 
138. Perng RP et al. (1997) A phase I feasibility and pharmacokinetic study of intrapleural 
paclitaxel in patients with malignant pleural effusions. Anticancer Drugs 8:565-73. 
139. Hughes RS (2005) Malignant pleural mesothelioma. Am}Med Sci329:29^4. 
140. Hagiwara A, Takahashi T, Ueda T, Iwamoto A and Torii T (1987) Activated carbon 
particles as anti-cancer drug carrier into regional lymph nodes. Anticancer Drug Des 
1:313-21. 
141. Hagiwara A et al. (1996) Methotrexate bound to carbon particles used for treating cancers 
with lymph node metastases in animal experiments and a clinical pilot study. 
Cancer 78:2199-209. 
142. Minato H et al. (1994) Survival of patients with gastric cancer treated with intra-lymph 
nodal injection of activated carbon particles absorbed mitomycin C. Gan To Kagaku 
Ryoho 21:2263-5. 
143. Natsugoe S et al. (1993) Loco-regional treatment for esophageal cancer with bleomycin 
adsorbed to activated carbon particles. Anticancer Res 13:1785-7. 
144. Hagiwara A et al. (1997) Selective drug delivery to peri-tumoral region and regional 
lymphatics by local injection of aclarubicin adsorbed on activated carbon particles in 
patients with breast cancer — a pilot study. Anticancer Drugs 8:666-70. 
145. Huang Y et al. (2002) Local injection of M-CH combined with i.p. hyperthermic hypoosmolar 
infusion is an effective therapy in advanced gastric cancer. Anticancer Drugs 
13:431-5. 
146. Edwards JM and Pheils PJ (1978) Endolymphatic isotope and BCG in the management 
of malignant melanoma. Aust N ZJ Surg 48:40-8. 
147. Vebersik V (1973) Direct endolymphatic therapy using radioisotopes. Strahlentherapie 
145:401-5. 
148. Chiappa S, Uslenghi C, Galli G, Ravasi G and Gonadonna G (1966) Lymphangiography 
and endolymphatic radiotherapy in testicular tumours. Br J Radiol 39:498-512. 
149. Dellepiane G, Tetti A and Davitti L (1965) Endolymphatic Radioisotope Therapy in 
Uterine Carcinomas. Minerva Med 56:2016-21. 
150. Stauch GW, Heissen E and Magnus L (1974) Lung-dosimetry within the framework 
of endolymphatic radionuclide therapy. Experimental and clinical results. Med Welt 
25:1036-8. 
Nanoparticles for Targeting Lymphatics 607 
151. Kenda R, Musumeci R and Uslenghi C (1975) Endolymphatic radiotherapy in 
malignant lymphomas: Its potential "prophylactic" value in cases with negative 
lymphograms. Lymphology 8:84-90. 
152. Wisner ER, Theon A, Griffey SM and Mclntire GL (2000) Long-term effect of irradiation 
on lymph node uptake of interstitially delivered nanoparticulate contrast media. Invest 
Radiol 35:199-204. 
153. Wisner ER, Theon AP, Katzberg RW, Griffey SM and Mclntire GL (1999) Lymph node 
uptake of interstitially delivered particulate contrast media before and after irradiation 
in dogs. Acad Radiol 6:119-25. 
154. Currier MA, Adams LC, Mahller YY and Cripe TP (2005) Widespread intratumoral 
virus distribution with fractionated injection enables local control of large human rhabdomyosarcoma 
xenografts by oncolytic herpes simplex viruses. Cancer Gene Ther. 
155. Duncan IC, Fourie PA and Alberts AS (2004) Direct percutaneous intratumoral 
bleomycin injection for palliative treatment of impending quadriplegia. AJNR Am } 
Neuroradiol 25:1121-3. 
156. Duvillard C, Romanet P, Cosmidis A, Beaudouin N and Chauffert B (2004) Phase 2 
study of intratumoral cisplatin and epinephrine treatment for locally recurrent head 
and neck tumors. Ann Otol Rhinol Laryngol 113:229-33. 
157. Hilger I et al. (2002) Thermal ablation of tumors using magnetic nanoparticles: An in vivo 
feasibility study. Invest Radiol 37:580-6. 
158. Gopalan B et al. (2004) Nanoparticle based systemic gene therapy for lung cancer: 
Molecular mechanisms and strategies to suppress nanoparticle-mediated inflammatory 
response. Technol Cancer Res Treat 3:647-57. 
159. Bao A, Goins B, Klipper R, Negrete G and Phillips WT (2003) 186Re-liposome labeling 
using 186Re-SNS/S complexes: In vitro stability, imaging, and biodistribution in rats. 
JNucl Med 44:1992-9. 
160. Ueno NT et al. (2002) Systemic gene therapy in human xenograft tumor models by 
liposomal delivery of the E1A gene. Cancer Res 62:6712-6. 
161. Villaret D et al. (2002) A multicenter phase II study of tgDCC-ElA for the intratumoral 
treatment of patients with recurrent head and neck squamous cell carcinoma. Head Neck 
24:661-9. 
162. Nishikawa M and Hashida M (1999) Pharmacokinetics of anticancer drugs, plasmid 
DNA, and their delivery systems in tissue-isolated perfused tumors. Adv Drug Del Rev 
40:19-37. 
163. Nomura T et al. (1997) Intratumoral pharmacokinetics and in vivo gene expression of 
naked plasmid DNA and its cationic liposome complexes after direct gene transfer. 
Cancer Res 57:2681-6. 
164. Bao A et al. (2005) Theoretical study of the influence of a heterogeneous activity distribution 
on intratumoral absorbed dose distribution. Med Phys 32:200-8. 
165. Emfietzoglou D, Kostarelos K and Sgouros G (2001) An analytic dosimetry study 
for the use of radionuclide-liposome conjugates in internal radiotherapy. / Nucl Med 
42:499-504. 
608 Phillips 
166. Kostarelos K and Emfietzoglou D (2000) Tissue dosimetry of liposome-radionuclide 
complexes for internal radiotherapy: Toward liposome-targeted therapeutic radiopharmaceuticals. 
Anticancer Res 20:3339—15. 
167. Kostarelos K et al. (2004) Binding and interstitial penetration of liposomes within avascular 
tumor spheroids. Int J Cancer 112:713-21. 
168. Tsyb AF, Drozdovskii B, Ikonnikov AI and Mukhamedzhanov I (1991) Intralymphatic 
administration of open radionuclides in complex treatment of rheumatoid arthritis. 
Med Radiol (Mosk) 36:12-5. 
169. van der Zant FM et al. (2004) Radiation synovectomy of the ankle with 75 MBq colloidal 
186rhenium-sulfide: Effect, leakage, and radiation considerations. / Rheumatol 
31:896-901. 
170. Galvao MM, lanhez LE and Sabbaga E (1982) Endolymphatic irradiation. A useful 
method for immunosuppression in renal transplantation. AMB Rev Assoc Med Bras 
28:55-8. 
171. Galvao MM, Peixinho ZF, Mendes NF, lanhez LE and Sabbaga E (1998) Endolymphatic 
irradiation in preparation for renal transplantation: A 26-year's follow-up. Sao Paulo 
Med /116:1710-4. 
26 
Polymeric Nanoparticles 
for Delivery in the Gastro-lntestinal 
Tract 
Mayank D. Bhavsar, Dinesh B. Shenoy 
and Mansoor M. Amiji 
1. Oral Drug Delivery 
In the last few decades, there has been a tremendous explosion in the research pertaining 
to novel (or advanced) drug delivery systems. Majority of the efforts have 
been directed towards development of "better" formulations of existing and/or 
off-patent drugs; the betterment being mostly aimed at improving the performance 
of the drug by altering the disposition and pharmacokinetics. Similarly, the trend is 
also being extrapolated to novel therapeutic compounds that are still in the pipeline, 
with the additional objective of positioning the molecule in highly competitive, 
technology-based intellectual property environment. The outcome has been phenomenal 
and the market size of advanced drug delivery systems is expected to 
swell to whopping US$ 40 billion by 2008, from its current size of US$ 20 billion. 
Non-invasive therapeutics has been the time-tested and most favored mode of 
drug administration. Oral route remains the front-runner in this segment. Current 
market share of the oral dosage forms is approximately 50% of all the formulations 
marketed and amounts to approximately US$ 40 billion. When a new chemical 
entity (NCE) is being developed, the first target of a formulation scientist would be 
to exploit the oral route. Often, a quick test to evaluate the oral bioavailability of the 
609 
610 Bhavsar, Shenoy & Amiji 
NCE is to fill the drug into hard gelatin capsules along with lactose, as this constitutes 
the simplest formulation that could be developed for oral administration. With 
the majority of novel drugs being highly hydrophobic or being of biotechnology 
origin, they pose serious and complicated challenges to the formulation scientists. 
Besides the ease of administration and patient compliance, the variety of excipients 
available (or being investigated) and the lesser cost involved for developing oral 
dosage forms, favor developments in this area of formulation science, compared 
with other delivery systems, especially those that involve invasive administration. 
The 21st century is being dedicated to nano-/bio-technological advancements 
and this has not spared the pharmaceutical product development section. "Nano" 
is the most widely used keyword that has penetrated almost every walk of life, 
and nanotechnology has become the key driving force behind the thriving hightechnology 
based pharmaceutical drug delivery industry. This chapter focuses on 
one of the components of widely explored product development showcase, that of 
polymeric nanoparticles. 
2. Anatomical and Physiological Considerations 
of Gastro-intestinal Tract (GIT) for Delivery 
To explore opportunities that are available for the delivery of bioactive compounds 
throughout the GIT, one has to first understand the anatomical and physiological 
conditions of the system because the secret of innovative formulation lies in exploiting 
these conditions as modulators for a well-programmed drug disposition. 
The human digestive system is specialized to perform functions such as ingestion, 
digestion and absorption. The organs of digestion are essentially divided into 
two main groups: the gastrointestinal tract or the alimentary canal and the auxiliary 
structures. The gastrointestinal tract is continuous tube-like structure beginning 
with the mouth (oral cavity) and extending further as pharynx, esophagus, 
stomach, small intestine, large intestine, rectum and finally culminating into the 
anal canal.1-6 The auxiliary structures include teeth, tongue, salivary glands, liver, 
gall bladder and pancreas. For the purpose of this chapter, our discussion will be 
limited to the anatomy and physiology of the gastrointestinal tract in its relation 
to oral drug delivery. Figure 1 provides a quick understanding of the GI targets, 
principles of formulation development that could be utilized and the application 
opportunities of the nanoparticles-based drug delivery system throughout the GIT. 
Different portions of the gastrointestinal tract serve different functions, but 
almost all the portions of the digestive tract are made up of four basic layers: 
(i) Mucosa, which is the mucus membrane, principally consisting of epithelial tissue 
and forming the inner most lining of the tract. In esophagus and anal canal, the 
mucus epithelium is specialized for protection of the underlying tissue. In other 
Polymeric Nanoparticles for Delivery in the Castro-Intestinal Tract 611 
Oral cavity 
Local or systemic 
therapeutics 
Principles of controlled 
delivery or quick onset or 
bioadhesion may be 
applied 
Applications: 
periodontitis, candidosis 
Small Intestine 
Mostly for systemic 
therapeutics 
Principles of controlled / 
delayed / sustained / 
pulsatile research maybe 
applied 
Possibility of utilization 
of stimuli-induced drug 
disposition 
Applications: gene 
delivery, vaccination, 
protein/peptide delivery, 
enhancement of 
bioavailability, 
controlled release 
Stomach 
Local or systemic 
therapeutics 
Principles of bioadhesion 
or gastric retention may 
be applied to enhance 
efficacy 
Applications: Gastric 
ulceration, mucosal 
immunization, gene 
delivery 
jf -V f 
\ f yi 
*J\ 
Larse Intestine 
Mostly for local 
therapeutics 
Principles of delayed 
release or triggered release 
may be used 
Applications: inflammatory 
bowel diseases 
Fig. 1. GIT targets, formulation principles, opportunities and applications. 
areas of the gastrointestinal tract, the epithelium is specialized for the secretion of 
mucus or digestive juices or for absorption, (ii) Submucosa, a thick layer of connective 
tissue containing nerves, blood vessels and glands, (iii) Muscularis, two layers 
of smooth muscles. The outer muscle layers are arranged longitudinally and the 
inner layer of muscles encircle the wall of the tract, (iv) Visceral peritoneum, the 
outermost layer of the tract and is a serous membrane, also known as serosa.1-6 
The mouth or the oral cavity comprises of the lips, cheeks, tongue, hard palate, 
soft palate and the floor of the mouth. The oral cavity is lined with the mucous membrane 
(oral mucosa) and includes the buccal, sublingual, gingival, palatal and labial 
mucosae.7-9 The oral mucosal surface has varied thickness with buccal mucosa having 
a thickness of 500-800 /xm, while the palates, gingivae and floor of the mouth 
measuring 100-200 /xm.10'11 The buccal and sublingual tissues are the principal 
focus for drug delivery via the oral cavity, because of the fact that they are more 
permeable than the other mucosal regions of the mouth. The oral mucosal surface 
comprises of less than 1% of the total surface area of the gastrointestinal tract, but is 
612 Bhavsar, Shenoy & Amiji 
high vascularized, allowing the drugs to diffuse from the oral mucosa and directly 
accessing the systemic circulation.8,9 Thus, the drugs entering the systemic circulation 
through the oral mucosa can bypass gastrointestinal tract and the first pass 
metabolism in liver. The permeability of the oral mucosa is greater for sublingual 
cavity, followed by buccal cavity and than the palatal surface. An enzymatic barrier 
also exists in the oral mucosa, which causes a rapid degradation of the peptides 
and proteins. The cells of the oral mucosa are surrounded by an environment of 
mucus, which is secreted by the mucous membrane, and is believed to play a role 
in the bioadhesion of mucoadhesive drug delivery systems.7-10,12 The pharynx and 
the esophagus also have the same anatomy and physiology as the rest of the gastrointestinal 
tract, but they are not generally considered as sites for drug delivery, 
and hence will not be discussed in this chapter. 
The esophagus ends into the stomach and is separated from the stomach by 
a cardiac sphincter muscle which acts as a valve system. It is a J-shaped, bag-like 
structure and described to have two curvatures, the concave curvature known as the 
lesser curvature and the convex curvature known as the greater curvature. Stomach 
is also essentially composed of the same four layers as the rest of gastrointestinal 
tract, which include the mucosal layer, submucosa, muscularis and serosa. The 
gastric mucosa contains many deep glands. These glands contain parietal cells 
which are responsible for the secretion of hydrochloric acid and the chief cells which 
secrete pepsinogens. Mucus is also secreted by these glands. The major barrier (or 
alternatively the opportunity) for drug delivery to the stomach is the low pH that 
exists in the organ because of the secretion of hydrochloric acid. The functions 
of stomach lie more in the digestion and it has very limited absorptive function. 
Delivery of proteins via the oral route faces a major hurdle in the stomach because of 
pepsinogens present in the gastric fluids which are responsible for the breakdown 
of proteins. 
The stomach ends into the small intestine and this region is guarded by a 
pyloric sphincter. The small intestine is further divided into three major parts which 
include the duodenum (25 cm in length), jejunum (2 meters long) and the ileum 
(3 meters long). The walls of the small intestine composed of the four layers previously 
described. The mucosa of the small intestine contains solitary lymph nodules 
and aggregated lymph nodules (Peyer's patches). Peyer's patches are found on the 
side opposite to the mesenteric wall of the intestine. They are usually oval in shape 
and occur more frequently in the distal areas of the small intestine and also at the 
terminal end of the colon. The Peyer's patches is comprised of four zones: (i) the germinal 
center which is in turn made up of three different cell types i.e. lymphocytes, 
macrophages and the dendritic reticular cells, (ii) small lymphocytic zone which 
shows the presence of lymphocytes and macrophages, (iii) interfollicular zone that 
is made up of lymphocytes which are loosely packed with large intercellular spaces, 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 613 
and (iv) subepithelial zone which shows large accumulation of macrophages and 
plasma cells. The Peyer's patches are lined above the lymphoid follicles by a membranous 
layer of epithelial cells called the follicle-associated epithelium (FAE). FAE 
is composed of absorptive cells, goblet cells, M cells and the enteroendocrine cells. 
These cells assist the Peyer's patches to transport macromolecules and particulate 
matter from the GIT into lymphatic/systemic circulation.13 In addition to the 
general structure, the small intestine also shows the presence of tiny finger-like 
structures known as villi, which are made of epithelial tissue overlying the blood 
and lymph capillary network. The free edges of the cells of the villi are divided 
into microvilli, which form the brush border. Throughout the length of the small 
intestine, the mucous membrane is covered by villi. The main function of small 
intestine is digestion and absorption of the food that is passed down from the 
stomach. The epithelial cells of the mucosa of the small intestine are specialized 
for the absorption of nutrients. The process of absorption in the small intestine is 
also assisted by its length and a very large surface area.14-16 The delivery of the 
drugs to the small intestine is preferred because drugs typically exhibit maximal 
absorption from this site, compared with other regions of the gastrointestinal tract. 
The absorption of drugs and particulate delivery system from the small intestine is 
believed to occur through gut associated lymphoid tissue and also from other nonlymphoid 
tissue.15 The mucus covers the mucosal layer of the small intestine that 
controls the absorption of nutrients, electrolytes and fluids, and also forms a physical 
barrier to the environment and absorption of drugs.14 The brush border enzymes 
form an enzymatic barrier for the absorption of proteins from the small intestine. 
The other major barrier to drug absorption to the small intestine is the action of 
ATP-dependent efflux protein P-glycoprotein pumps (PGPs), which exists on the 
cell membrane of the intestinal epithelium. PGPs transport certain drugs actively 
back into the intestinal lumen. PGPs are a part of the protective barrier of the small 
intestine that limits absorption of potentially toxic substances.16 
The small intestine ends into the large intestine. It is called the large intestine 
because of the larger diameter of the tract compared to the small intestine. It is 
approximately 1.5 m in length beginning at caecum and ending in rectum and the 
anal canal. There is a difference between the wall of the large intestine and small 
intestine. The large intestine shows absence of villi structure and contains simple 
columnar cells with numerous goblet cells. The goblet cells secrete mucus that 
lubricates the colonic content as it passes through the colon. The submucous layer 
of the large intestine consists of more lymphoid tissue than any other part of the 
alimentary canal to provide non-specific defense against invasion by microbes in 
the food and the bacterial flora that resides in the gut. Drug delivery to the large 
intestine via the oral route for local action is a challenging task, as the drug carrier 
system will have to face the rigors of the preceding sections of the GIT before 
614 Bhavsar, Shenoy & Amiji 
reaching the desired site of action. Rectal delivery of drugs is an alternative for 
local action, but it suffers the disadvantage of patient compliance. The mucus layer 
of the large intestine can take up particles in a particular size range and this property 
could be exploited for delivery of the drug to the large intestine.17'18 
3. Introduction to Polymeric Nanoparticles as Carriers 
Modern day drugs are very effective in treating disease, but many of these drugs 
have limitations when it comes to the route of administration. Major advances in the 
field of biochemistry and biotechnology have led to the findings of a large number of 
bioactive molecules and vaccines, which are based on peptides, proteins and nucleic 
acids. Oral route is the most desired for administration for all drugs and bioactive 
molecules, but some of these drugs and molecules cannot be administered orally 
due to the fact that they become inactive in the GIT before getting absorbed, mainly 
due to enzymatic degradation. Hence, the parenteral route of drug administration 
becomes a very effective route for dosing of such drugs. However, the parenteral 
route of drug administration has the problem of being inconvenient for self administration 
by the patient and hence reduces patient compliance.14,19-23 In the last few 
years, we have seen rapid development of drug delivery systems for the treatment 
of human diseases, which is the direct result of the extensive research being done 
on the applications of materials for medical and pharmaceutical product development. 
These advanced drug delivery systems include mostly colloidal carriers like 
liposomes, niosomes, nanoparticles, dendrimers, nanosuspensions, micelles and 
nano-/micro-emulsions.24-31 These drug carrier systems offer many advantages 
like improved efficacy, reduced toxicity and improved patient compliance, and 
are also cost effective in many cases over conventional drug delivery systems.32,33 
Among the above mentioned colloidal drug delivery systems, nanoparticles represent 
the most appealing therapeutic nanocarrier systems by comprehensively 
addressing majority of the issues like stability, scalability, reproducibility, and by 
offering the best compromise between the efficacy and applicability.14,26,34-40 
Nanoparticles can be defined as solid colloidal particles, produced by mechanical 
or chemical means, which are typically in the nanometric size range (1 to 
1000 nm).19,32,33 Nanoparticles, especially those prepared from polymeric materials, 
enjoy tremendous popularity due to ease of preparation, easy to tune the physicochemical 
properties (e.g. through an array of polymeric materials), possibility of 
surface modification, excellent stability, and scalability to industrial production. 
Since their conception in the mid 70s, nanoparticles have found applicability in 
almost every section of medicine and biology (besides host of other fields) in general, 
and also for controlled and / o r targeted delivery of drugs and genetic materials 
in particular.26,34,37-39,41-52 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 615 
The basis in the development of nanoparticles lies in Paul Elrich's idea of 
designing "magic bullet" carrying active molecules in them and being able to target 
specific sites in the body for the desired therapeutic effects.33 Depending on the 
process by which they are prepared, these systems can be classified as nanospheres 
(nanoparticles) having a dense and solid polymeric network (monolithic matrix), 
or as nanocapsules which consist of a hollow core surrounded by a polymeric 
shell.32'33 Drug-loaded nanoparticles have been developed for almost every route 
of administration, i.e. nasal, ocular, mucosal, inhalation, oral, transdermal and 
parenteral.14/37,41,53~56 Clinically, they have found applications for diagnosing and 
treating a wide range of pathological conditions. 
Nanoparticles can be prepared from both synthetic and natural. The polymeric 
materials could be either biodegradable or non-biodegradable, but should 
be essentially biocompatible. Poly (DL-lactide-co-glycolide) (PLGA), poly (ecaprolactone) 
(PCL), poly (alkylcyanoacrylates), poly (styrene-co-maleic anhydride), 
poly (divinylether-co-maleic anhydride), poly (vinyl alcohol), poly (ethylene 
glycol) are some examples of synthetic, non-immunogenic polymers extensively 
used for nanoparticle preparation. Similarly, poly (amino acids), albumin, gelatin, 
hyaluronic acid, dextran, starch, and chitosan are some of the natural biodegradable 
polymers. While each of polymers poses its own advantages and nanoparticles can 
be synthesized with high degree of reproducibility from a majority of them, natural 
polymers, due to their natural origin, have preference, considering non-toxicity 
and biodegradability. The striking advantage of synthetic polymers remains the 
possibility to synthesize them reproducibly with well-defined physico-chemical 
properties. Advancement in biotechnology is helping the natural polymers to 
overcome this drawback and we can expect a surge in delivery systems based 
on them. 
Polymeric nanoparticles have been extensively researched for their applicability 
as oral drug carrier systems. In the following sections, we will discuss how 
they are being explored in buccal cavity therapeutics, as stomach specific delivery 
systems, for mucosal targeting in the small intestine, and for the treatment of 
inflammatory bowel disease. 
4. Preparation of Polymeric Nanoparticles 
There are several methods on the preparation of polymeric nanoparticles and incorporation 
of bioactive compounds into them. In general, one of the two principles 
methods is utilized: controlled precipitation or controlled dispersion of the 
polymer. Few of the popular methods include solvent displacement, salting-out, 
emulsion-solvent-evaporation, emulsion-solvent-diffusion, polymerization, complexation 
and supercritical fluid technology. Figure 2 provides an overview of 
Solvent 
Displacement 
Salting-out Emulsion-solvent 
diffusion 
Polymer + Drug in 
water miscible 
organic solvent 
Emulsion-solvent 
evaporation 
Supercritical fluid 
technology 
Polymer + Drug in water 
miscible organic solvent 
Polymer + Drug in partly 
water miscible organic solvent 
Added lo 
V j 
1' 
Non solvent aqueous phase 
+ stabilizer + stirring 
Aqueous gel + 
salting out agent + stabilizer 
Polymer + Drug in 
organic /aqueous 
solvent 
Polymer + Drug 
in organic solvent 
Sprayed via no«le 
Formation of polymeric 
nanoparticles 
Fig. 2. Schematic representation of methods used for preparation of p 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 61 7 
the methodologies and technologies available for the preparation of polymeric 
nanoparticles. 
In the case of solvent displacement method, which is the simplest of all, the polymer 
is dissolved in a good solvent that maybe partially-polar and water-miscible 
solvent such as ethanol or acetone.33 When the drug is to be incorporated into the 
particles, it can be dissolved in the same phase along with the polymer. This polymer 
phase is introduced into a non-solvent aqueous phase containing a stabilizer 
(generally a hydrophilic surfactant) at a controlled rate under continuous mixing. 
As the partially-polar solvent diffuses rapidly into the aqueous phase (i.e. as the 
partially-polar phase is displaced by the polar phase), the polymer starts precipitating 
due to changes in its solubility, resulting in the formation of nanoparticles. 
The surfactant present in the aqueous phase helps in preventing particle aggregation. 
Choice of a drug/polymer/solvent/non-solvent system is the major limitation 
of this method and hence its applicability is confined to hydrophobic drugs and 
polymers. 
Salting out technique is generally used for the preparation of drug-loaded 
biodegradable nanoparticles. This method was first applied to pseudolatexes.33 
It is based on the separation of water-miscible solvent from aqueous solutions by a 
salting out effect. An o /w emulsion is formed by adding a solution of the polymer 
and the drug in a water miscible solvent into an aqueous gel containing a salting-out 
agent and a colloidal stabilizer. Water is added to dilute this mixture, as a result of 
which nanoparticles are formed. Solvent and salting-out agents are then removed 
by cross-flow filtration. The use of this method results in a very high loading efficiency 
along with high yield and also the scale-up is fairly easy, but this method 
can only be used for the loading of lipophilic drugs. 
Emulsification-solvent-evaporation is based on the formation of a biphasic 
(o/w or w/o) or triphasic ( w / o /w or o/w/o) emulsion.33 Generally, a preformed 
polymer is dissolved in an organic solvent which is water immiscible along with 
the drug, and is emulsified in an aqueous solution (o/w emulsion). The formed 
emulsion is then exposed to high energy mixers (e.g. high-speed or high-pressure 
homogenizers, colloidal mills or ultra sonic devices) to reduce globule size. The 
organic solvent is removed either by using heat or vacuum or even both at times. 
Nanoparticles are obtained as fine aqueous dispersions which can be collected and 
purified. The process variables involved in this method are complex and manifold, 
and the nanoparticles obtained are often polydisperse. However, this method is 
very popular for preparing polymeric microparticles rather than nanoparticles, as 
it facilitates industrial applicability and scalability. 
Emulsion-solvent-diffusion method is another method which is used for 
nanoparticles preparation. It is a modified salting-out technique and differs mainly 
in the organic solvent which is partially miscible with water in this case.32 This 
618 Bhavsar, Shenoy & Amiji 
solvent is pre-saturated with water to achieve initial thermodynamic equilibrium 
between water and the organic phase. Solvent diffuses out upon addition of water 
and results in the formation of nanoparticle suspension. 
Controlled complexation induced by electrostatic interactions between oppositely 
charged polymers can yield stable colloidal dispersions. The interacting polymers 
could be therapeutically active (e.g. oligonucleotides and plasmid DNA) or 
may have tailored properties (e.g. pH-sensitivity).57,58 A wide variety of chargebearing 
polymers can be utilized to manufacture composite nanoparticles and 
varying physico-chemical properties.48,59-63 
Supercritical fluid technology is an emerging science for the production of micro 
and nanoparticles.64,65 In this method, an organic liquid solution of the polymer and 
the active moiety is sprayed through a nozzle into a chamber containing a gas that 
is miscible with the solvent, but in which the polymer and the active compound 
are not soluble. The gaseous phase in this case is a super critical fluid (e.g. supercritical 
C02). The dispersion of the liquid solution in such a condition generates a 
high degree of super-saturation, leading to the formation of fine, uniform colloidal 
particles. The particles can be recovered from the solution by depressurizing the 
chamber and allowing the gas to escape.66,67 
While all of the above mentioned procedures employ preformed and well characterized 
polymers, there are other techniques for obtaining fine nanoparticles from 
monomers via in situ polymerization pathway. The most popular example for this 
method of synthesis is the nanoparticles made from poly (methylmethacrylates), 
poly (alkylcyanoacrylates) and poly (methylidenemalonates).68 Generally, a water 
insoluble monomer is dispersed in an aqueous medium containing a colloidal stabilizer, 
and the polymerization is induced and controlled by the addition of a chemical 
initiator or by variations in physical parameters such as pH or radiation. Both 
hydrophilic and lipophilic drugs can be entrapped in the polymeric wall when 
added to the polymerization medium or adsorbed on preformed particles. 
While each of the above mentioned nanoparticle preparation method has its 
advantages and disadvantages, they can all be fine-tuned to encapsulate variety of 
drugs. The literature evidence shows that the nanoparticles are mostly employed to 
incorporate hydrophobic drugs, simply because the majority of the techniques facilitate 
encapsulation of lipophilic compounds with very high loading (approximately 
up to 40% by weight) and capturing efficiencies (nearly 100%). When hydrophilic 
drugs are to be incorporated, in situ polymerization or complexation remains the 
most accepted method. 
The collective advancements in nanotechnology and engineering sciences are 
expected to contribute major breakthroughs for bulk manufacturing of polymeric 
nanoparticles. In the highly competitive pharma/biotech industry, the formulation 
scientist can concentrate towards development of novel products, irrespective of 
Polymeric Nanoparticles for Delivery in the Castro-Intestinal Tract 619 
complexities involved in the procedures. As in most cases, majority of the scaleup 
issues can be addressed and solved with the help of parallel advancements in 
high-technology engineering. 
5. Design Consideration for Nanoparticle-based 
Delivery Systems 
Polymeric nanoparticles, since their first appearance in the 70s, have been keenly 
explored as delivery systems for small drug molecules, and also for macromolecules 
like nucleic acids, proteins, hormones and peptides.69'70 With the patent protection 
to a number of blockbuster drugs expiring in this decade, innovative dosage forms, 
such as polymeric nanoparticles, can form a very powerful drug delivery technology 
for the pharmaceutical industry. Such technology-based products can be used 
for the extension of the patent life of the drug, or to prevent /delay the entry of 
generic versions into the specialized markets. In general, the polymeric drug delivery 
systems offer advantages such as the reduction in the total dose (hence also the 
dosing frequency), reduced side effects, delivery with enhanced efficiency (hence 
better performance), and most importantly, improved patient compliance.70 
When designing a polymer-based nanoparticulate drug delivery system, the 
choice of synthetic or biopolymer is the most important consideration. In the following 
section, we will discuss different characteristics of polymers that play an 
important role in the final product and hence should be considered during preformulation 
stages. 
5.1. Polymer characteristics 
Polymers have become a vital and integral part for the development of any novel 
drug delivery system. Factors such as chemical structure of the polymer, composition, 
molecular weight, morphology of the polymer (amorphous/crystalline/ 
residual stresses), size of the delivery system, process of degradation (enzymatic or 
non-enzymatic) etc. govern the behavior of polymer within the body69 The polymer 
selection itself is a judicious process and following factors should be considered 
while making a decision: 
Basic requirements: 
• Non-reactive: Chemical inertness with respect to active compound and the 
biological environment. 
• Biocompatibility: Should be compatible with living cells and tissues that come in 
contact with the polymer. 
• Non-pyrogenic: Should be free of any pyrogenic factors. 
620 Bhavsar, Shenoy & Amiji 
• Impurities: All the impurities should be well-established and present in minimal 
amounts. The impurities, if present, should also be biocompatible or should not 
pose any toxicity at amounts present. 
• By-products: If the polymer undergoes any kind of biotransformation upon introduction 
into the body, the by-products should also be biocompatible. 
• Regulatory issues: The polymer should be available in the cGMP grade and must 
be approved by the regulatory authorities for human use. 
Specialized requirements: 
• Loading capacity: If complexation or chemical conjugation is the method used for 
preparation, then the polymer must have sufficient reactive groups to promote 
respective interactions. 
• Permeability: Permeability to molecules and water will govern the diffusivity 
and the release of the payload. 
• Swellability: This could be of relevance when designing a floating or a bioadhesive 
system 
• Viscoelasticity: Could be an important controlling parameter for gel-forming and 
adhesive systems. 
• Sensitivity to environment: The triggering factor could be the pH, specific 
enzymes, or even the microbial flora prevailing in the GIT. 
While the maneuverability around each of the above factors is very limited for 
a sterile dosage form, it becomes more flexible while developing an oral product. 
Parallel developments in the field of excipient science have contributed to a range of 
new high performance polymers, making the choice of an approved polymer easier 
for the formulation scientist. We recommend that interested readers should visit 
the websites of major manufacturers of pharmaceutical excipients (e.g. Eastman 
Chemicals, FMC Biopolymer, Colorcon, Gattefosse, Croda, Lipoid, Noveon, BASF, 
Roehm Pharma, Degussa, etc.) to build-up a database. 
5.2. Drug characteristics 
The performance of the dosage form for any bioactive compound will depend on 
physico-chemical properties of the drug, as well as the auxiliary factors. Few of 
the factors belonging to the former are molecular weight, solubility (aqueous or 
organic), partition coefficient, crystallinity and ionic properties. As these are the 
inherent properties of the drug, there is less scope for tailoring them to manipulate 
in vivo behavior of the formulation. However, factors such as solubility, for example, 
can be tuned to a certain extent by altering the particle size or using certain excipients 
(e.g. cyclodextrins). Some of the auxiliary factors that should be considered 
while designing a nanoparticle-based oral system are the dose of the drug, site of 
Polymeric Nanoparticles for Delivery in the Castro-Intestinal Tract 621 
action of the drug (e.g. absorption is limited to certain segments of the GIT only), 
stability of the drug (both under in vitro and in vivo conditions) and the desired 
pharmacokinetics/distribution profile. 
5.3. Application characteristics 
In the majority of the cases, the formulation parameters are decided based on 
its intended application. Upon oral administration, the nanoparticles could be 
expected to exert local action along the GIT, or could be used to deliver drugs 
to the systemic compartment, or could be meant for uptake by cells lining the 
GIT. The duration and kind of desired pharmacological action can be used as the 
guiding principle in designing the delivery module (fast versus sustained; regular 
versus controlled; unstimulated versus triggered; continuous versus pulsatile). 
More of these principles of design have been discussed under following section 
citing appropriate examples. 
6. Nanoparticles in Experimental and Clinical Medicine 
Many of the principles of drug design and delivery are based on naturally occurring 
phenomena (bio-mimicking approach) and the same principles guide the applicability 
of nanocarriers in therapeutics. The striking advantage of the nanoparticles 
is the large surface area that they offer when presented in a biological environment 
and the flexibility to alter the physico-chemical properties by manipulating the 
core polymer or by surface nanoengineering. Many clinical situations and conditions 
demand specialized therapeutics to achieve improved level of healing. In such 
situations, the requirements are specified by the clinician, which form the basis of 
product development. Table 1 gives an overview of the applicability of polymeric 
nanoparticles via oral route. 
6.1. Drug delivery in the oral cavity 
Buccal cavity mucosa has been studied for polymeric drug delivery. Bioadhesive 
or mucoadhesive (used when adhesion is to the mucosal tissue) polymers have 
been extensively used in buccal drug delivery because of the fact that a major 
limitation of the buccal cavity is the lack of dosage form retention. Polymers 
such as poly (methacrylate) derivatives, cyanoacrylates, epoxy resins, polystyrene, 
polyurethanes, hydroxypropyl methylcellulose, chitosan and poly acrylic acid have 
been studied for their potential use in buccal cavity therapeutics.7'8'71-81 
Poly (propylcyanoacrylate) (PPC A) nanoparticles have been studied as a potential 
carrier for the prophylactic treatment of candidosis.71 Candida albicans is a common 
organism which is found in the oral cavity. It occurs in the commensal form 
622 Bhavsar, Shenoy & Amiji 
Table 1 Summary of polymeric nanoparticle-based delivery systems for in the GIT. 
Site of action Incorporated 
compound 
Oral cavity FITC 
N/A 
Stomach Carbazole 
Amoxicillin 
pCMV-lacZ 
Small intestine Streptomycin 
Theophylline 
in depot tablets 
Tetanus toxoid 
Indomethacin 
Carbon-14 
5-Fluoroudine 
RBITC 
Vancomycin 
Iodine-125 
Valproic acid 
pCMV-lacZ 
N/A 
Phenobarbital 
Amifostine 
H. pylori lysate 
DNA 
pCR3Arah2 
mEpo gene 
Rifampicin 
Pyrazinamide 
Ketoprofen 
Isoniazid 
Calcitonin 
Heparin 
Polymer employed 
Poly(propylcyanoacrylate) 
Lectin-Gliadin 
Gliadin 
Gliadin 
PLGA 
Chitosan 
PLGA 
Poly(ethylene 
glycol-Poly(lactic acid 
PLGA 
Poly(methyl methacrylate) 
(PMMA) 
Poly(methylvinylether-comaleic 
anhydride) 
Poly(methylvinylether-comaleic 
anhydride) 
PLGA 
Polystyrene 
PLGA 
Poly(ethylene 
oxide)-poly(propylene 
oxide) 
Sulfobutylated-poly(vinyl 
alcohol)-PLGA 
PLGA 
PLGA 
PLGA 
Chitosan 
Chitosan 
Chitosan 
Lectin-PLGA 
Lectin-PLGA 
PLGA 
Lectin-PLGA 
Poly(N-isopropylacrylamide) 
Poly(N-vinylacetamide) 
Poly(t-butyl methacrylate) 
PLGA 
PCL 
PLGA 
Eudragit® RS and RL 
Size 
(nm) 
100-900 
500-600 
400-500 
250^00 
>200 
50-500 
200-260 
150-170 
100-200 
100-160 
200-250 
200-300 
100-200 
50-3000 
100-200 
150-190 
100-130 
100-200 
200-300 
300^00 
50-75 
100-1000 
70-150 
300^00 
300^:00 
100-200 
300-400 
148-895 
148-896 
148-897 
200-400 
270-300 
250-270 
250-280 
Reference 
71 
86 
130 
85 
87 
19 
131 
132 
139 
91 
136 
138 
139 
90 
139 
113 
14 
139 
143 
116 
112 
114 
111 
118 
118 
139 
118 
11, 95,134 
11, 95,134 
11, 95,134 
135 
137 
137 
137 
Polymeric Nanoparticles for Delivery in the Castro-Intestinal Tract 623 
Table 1 (Continued) 
Site of action Incorporated 
compound 
Insulin 
CyA 
Fluorescein 
Polymer employed 
Poly(iso-butyl cyanoacrylate) 
Polyesters 
Poly(methacrylic 
acid)-g-poly(ethylene glycol) 
Poly(alkylcyanoacrylate) 
Chitosan 
Poly(isobutylcyanoacrylate) 
Poly(acrylicacid)-gpoly(
ethylene glycol) 
PLGA 
Poly(fumaric-co-sebacic) 
anhydride 
PLGA 
Poly(methacrylic acid 
methacrylate) 
Hydroxypropyl 
methylcellulose 
Chitosan 
Gelatin 
PLGA 
PCL 
Eudragit® RS and RL 
Polystyrene 
Polystyrene nanoparticles 
coated 
with poloxamer 188 and 407 
Large intestine Fluorescent dye Polystyrene 
Rolipram PLGA 
Size 
(nm) 
85 
300-500 
200-1000 
1000^00 
200^00 
100-500 
200-100 
>1000 
>1000 
100-200 
30-110 
50-60 
150 
140 
100-200 
100-130 
170-310 
50-3000 
60 
100-1000 
300-500 
Reference 
142 
104 
103 
100 
97 
99 
103 
102 
102 
139 
22 
108 
106 
106 
139 
140 
141 
16 
133 
17 
18,123 
in healthy individuals, but it can become pathogenic to the body under conditions 
such as cancer chemotherapy, diabetes mellitus and also during antimicrobial therapy. 
The first step in candidosis infection is the adherence of the microorganisms 
to the epithelial cells of the host. The basic motivation of the investigation was to 
evaluate the ability of PPC A nanoparticles to disrupt the process of adherence of the 
microorganism onto the host cells. PPC A nanoparticles were prepared by emulsionpolymerization 
from propylcyanoacrylate monomers. Different kinds of surfactants 
were used for stabilization of nanoparticles, which resulted in the formation 
of nanoparticles having different size ranges. Surfactants like Tween® 80, Pluronic® 
P123, Tetronic® 904, docusate sodium, sodium oleate and sodium laurylsulfate 
produced particles in the nanometer range, whereas cetrimide, benzalkonium 
chloride and cetylpyrimidine chloride produced particles in the micrometer 
624 Bhavsar, Shenoy & Amiji 
range. Tetronic® 904 produced the smallest particles of the size 90 ± 10 nm. 
C. albicans blastospores were treated with the nanoparticle suspension and these 
treated blastospores were exposed to the buccal epithelial cells to check for adherence. 
It was found that nanoparticle treated blastospore adherence per buccal 
epithelial cells was reduced by up to 73%. The findings of this study may offer the 
basis for a prophylactic treatment of candidosis in immuno-compromised patients. 
Periodontal diseases are one of the major causes of teeth loss and it includes a 
number of diseases involving the supporting tissue of the teeth. Conventional methods 
of treatment of periodontitis include periodontal surgery and chemotherapy, 
but both these treatments cannot prevent the reoccurrence of the disease. Recently, 
a method for treating periodontitis, by using polymeric nanoparticles loaded with 
photosensitizer compounds, has been proposed.82 The nanoparticles exhibit controlled 
release of the photosensitizer molecule through the matrix polymer. The 
proposed application uses photosensitizer molecules such as porphyrins, chlorines, 
pheophorbides, bacteriopheophorbides, phthalocyanines, naphthalocyanines, thiazines, 
xanthenes, pyrrylium dyes, psoralens, quinones and amenolevulic acids. 
These compounds are either incorporated or complexed with nanoparticles made 
from biodegradable or non-biodegradable polymers. The effectiveness of photosensitizers 
relies on their association with cellular membranes, thereby targeting 
highly sensitive membranous intracellular organelles that control critical metabolic 
functions. The hydrophobic character of the photosensitizers means that they cannot 
be administered directly to a hydrophilic environment due to a tendency to 
aggregate (by molecular stacking, precipitation or other mechanisms), which can 
severely curtail photosensitization processes. Thus, they require formulation in carriers 
which are able to provide a hydrophobic environment to maintain them in a 
non aggregated form in both the formulation and in aqueous preparations prior 
to use83. These nanoparticles can then be applied by the dentist to the periodontal 
pockets in the form of gel which hardens on application. This allows for a slow 
and extended period of release, which can be fine tuned by choosing biodegradable 
or non-biodegradable polymer of the photosensitizer from the nanoparticles, 
and hence do not affect the normal cell function. The use of nanoparticles prevents 
the degradation of the photosensitizer molecule in the presence of saliva, 
white blood cells and other natural defenses in the mouth. Higher concentrations 
of the photosensitizer allows for a more effective treatment and this can be achieved 
by using specialized nanoparticles formed out of dendrimer-photosensitizer complexes. 
Furthermore, many of the dental diseases are difficult to treat due to a lack 
of accessibility and quick flushing of the dosage form by the saliva. The nanoparticles 
can be especially useful in such situations. The size of the carriers enables them 
not only to reach deeper parts of the infected area, but also to be retained at the site 
of action. 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 625 
6.2. Gastric mucosa as a target for oral 
nanoparticle-mediated therapy 
The major function of the stomach is to digest food and pass down the chyme to the 
intestine. The principal hurdle to the successful delivery of active compounds to 
the gastric mucosa using conventional delivery system is the gastric emptying time. 
These conventional delivery systems do not remain in the stomach for prolonged 
periods due to their inability to deliver the drug to the desired site in effective 
concentration and in fully active form. The other barrier to the delivery of drug is 
the mucus layer of the gastric mucosa. The primary component of mucus is glycoprotein 
which forms a dense condensed and complex microstructure, by forming 
numerous covalent and non covalent bonds with other mucin molecules. 
Helicobacter pylori has been recognized as a major gastric pathogen responsible 
for a variety of clinical manifestation including the development of gastritis, gastric 
ulcer and gastric carcinoma.84 It is a gram negative, spiral, urease producing 
microorganism isolated by Warren and Marshall in 1982. Umamaheshwari et ah,85 
studied the effectiveness of mucoadhesive nanoparticles bearing amoxicillin for the 
treatment of H. pylori. Mucoadhesive nanoparticles prepared from gliadin, having 
a size range of 285 to 392 nm, were used in the study. Gliadin is a group of polymorphic 
proteins extracted from gluten and are soluble in ethanolic solutions. They 
have a very low solubility in water except at extreme pH.86 In vitro stability study of 
gliadin and amoxicillin was performed in simulated gastric fluid and confirmed by 
HPLC. In vivo mucoadhesion capacity was evaluated by oral administration of fluorescent 
labeled gliadin nanoparticles. Size dependent mucoadhesive propensity 
and specificity was exhibited by gliadin nanoparticles with less than 300 nm particles 
showing 68% mucoadhesion, and more than 300 nm particle showing 75% and 
above mucoadhesion. Amoxicillin loaded gliadin nanoparticles were administered 
to Mongolian gerbils previously inoculated with human H. pylori to study in vivo 
clearance time (4 hrs, 8 hrs and 12 hrs), and placebo gliadin nanoparticles were also 
used as a control. Although amoxicillin loaded nanoparticles showed 100% inhibition 
of H. pyloi within 4 hrs of administration, it could not completely eradicate the 
H. pyroli in vivo. This study showed that amoxicillin loaded nanoparticles exhibited 
a longer gastric residence time than conventional amoxicillin formulation and also 
that topical action of amoxicillin on the gastric mucosa plays an important role in 
the clearance of the bacterium. 
Gastric mucosa can also be explored for the delivery of genetic material or for 
vaccination. A recent investigation explored PLGA nanoparticle stabilized with a 
cationic surfactant (dimethyldioctyldecylammonium bromide) as gene carriers for 
transport through the gastric mucosal barrier.87 Composite polymeric nanoparticles 
having a magnetic element and loaded with anti-metabolites have also been 
626 Bhavsar, Shenoy & Amiji 
explored for the treatment of gastric tumors.88 The magnetic component helps in 
external guiding and localization of the nanoparticles at the site of action. 
6.3. Nanoparticles for delivery of drugs and vaccines in the 
small intestine 
Gastrointestinal tract provides a variety of barriers, including proteolytic enzymes 
in gut lumen and on the brush border membrane, mucus layer, gut flora and epithelial 
cell lining, to the delivery of drugs. Factors which govern the uptake of particles 
from the gut include particle size, physico-chemical nature of particles, surface 
charge and attachment of uptake enhancers such as lectins or poloxamer. After 
oral administration of nanoparticles, they could be (i) directly eliminated in the 
faeces, (ii) adhering to the cells (bioadhesion) and /or, (iii) undergo oral absorption 
as a whole. Oral absorption of the nanoparticles results in passage across the gastrointestinal 
barriers and delivery of the payload into the blood, lymph and other 
tissues. Before this translocation can occur, the nanoparticles have to adhere to the 
surface of the intestine. Translocation of particles across the gastrointestinal wall 
can occur due to intracellular uptake by the absorptive cells of the intestine or paracellular 
uptake (i.e. between the cells of the intestinal wall), or phagocytic uptake 
by intestinal macrophages, or uptake by the M cells of the Peyer's patches.89 
Jani et a/.90 have shown that particle size plays a major role in the uptake of 
particles. They measured uptake by using radiolabeled polystyrene nanoparticles 
ranging from 50 nm to 3.0 /xm. They have been able to show that lower size particles 
(50 nm particles showed a 12% uptake by the cells of the small intestine) are 
taken up at a higher rate by the small intestine when compared to the larger particles 
(1 fim particles showed only 1% uptake by the cells of the small intestine). 
The lower size particles (<500nm) were detected in blood after intestinal uptake 
whereas larger size particles (>500nm) where not detected in blood. Also, these 
nanoparticles were detected in other tissues such as liver and spleen. A low surface 
charge on the surface of nanoparticles is desirous for good absorption. While 
Pluronic® or poloxamer (188 and 407) coating onto the surface of 50 nm polystyrene 
nanoparticles inhibited uptake in the small intestine, a similar coating on the 500 nm 
polystyrene nanoparticles showed an increased intestinal uptake. 
There has been yet another report to study the effect of surface modification on 
the uptake of polymeric nanoparticles using 14C-labeled poly (methylmethacrylate) 
(PMMA), having a mean size of 130 nm and coated with polysorbate (Tween®) 80 
or poloxamine 908.91 These nanoparticles were administered orally to rats and they 
were checked for their organ distribution. High radioactivity levels were observed 
in the stomach contents, below 5% radioactivity was detected in the stomach wall 
for the coated particles. Highest amount of radioactivity (about 40%) was found 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 627 
in the small intestine, confirming that these coated particles were absorbed in the 
small intestine. 
Developments in the field of polymer science have made the delivery of proteins 
and peptide drugs via the oral route possible, by protecting these molecules 
against pH/enzyme-induced degradation and also by prolonging the time of delivery 
to the mucosal sites.23'92"95 The most popular peptide used for oral delivery 
using polymeric nanoparticles is insulin. The first attempt to deliver insulin via 
the oral route was made by Couvreur et al.92 in 1980. Insulin was adsorbed on the 
surface of 200 nm poly (alkylcyanoacrylate) nanoparticles and administered orally 
to diabetic rats to seek hypoglycemic effects. The investigators did not observe 
any decrease in glucose level upon oral administration, but good hypoglycemic 
activity was observed upon subcutaneous administration, suggesting that insulin 
was getting degraded in the GIT. In another investigation, nanoparticles made out 
of poly (isobutylcyanoacrylate) (PIBCA) loaded with insulin when administered 
orally, resulted in a 50-60% reduction in the blood glucose levels of diabetic rats.96 
The onset of action was after 2 days of administration, but was seen for 20 days 
depending on the insulin dose. These results suggested that PIBCA nanoparticles 
successfully protected insulin against degradation in the GIT.93'97'98 A publication 
by the same group reported the ability of PIBCA nanoparticles to protect insulin 
from degradation by proteolytic enzymes, and thus providing nanoparticles based 
formulation for biologically active insulin for oral administration.99 Insulin labeled 
with Texas Red® was used for release studies and microcopy observations. The 
results obtained from fluorescence and confocal microscopy revealed the presence 
of concentrated fluorescent spots into the mucosa and even in the lamina propria. 
This suggested that these nanoparticles could cross the barrier presented by the 
intestinal epithelium. 
A patent was issued in 1995 for controlled release of insulin from biodegradable 
nanoparticles.100 Insulin was complexed with different polycyanoacrylate 
monomers at low pH and nanoparticles were prepared from this complex by anionic 
polymerization process. These nanoparticles were dosed orally to rats and blood 
glucose levels were monitored over four hours. A considerable decrease in blood 
glucose levels was observed in a group dosed with insulin loaded nanoparticles, 
compared with the untreated group. More recently, Pan et al.97 studied the effects 
of bioadhesive chitosan nanoparticles for improving the intestinal absorption of 
insulin in diabetic rats. Chitosan was chosen as the polymer for preparing the delivery 
system, because it exhibits strong electrostatic interaction with insulin, hence 
improving the loading efficiency of the polymer. It was also used for its bioadhesive 
properties for prolonged stay in the gastrointestinal tract, which in turn resulted in 
prolonged release times for insulin.101 A dose dependent decrease in blood glucose 
levels was observed after oral administration of these 290 nm particles in diabetic 
628 Bhavsar, Shenoy & Amiji 
rats. Chitosan-insulin nanoparticles showed a higher decrease in blood insulin 
levels when compared with chitosan-insulin solution, suggesting that they could 
enhance the intestinal absorption of insulin by promoting protection from gastric 
clearance, and also rendering longer resident time in circulation. 
Biodegradable polymers like PLGA, poly lactic acid (PLA), and poly (fumaric 
anhydride-co-sebacic anhydride) have been explored for the preparation of 
nanoparticulate formulations of insulin.102 Although the major finding of the study 
was intact bioactivity of insulin after intraperitoneal injection, it was also indicated 
that the nanoparticles prepared in the presence of Fe304 showed the best hypoglycemic 
results, and were also proved to be orally effective. 
Foss et «/.103 developed nanospheres from methacrylic acid grafted with poly 
(ethylene glycol) and also acrylic acid grafted with poly (ethylene glycol) as oral 
insulin carriers. From the results obtained after oral administration, it can be learned 
that diabetic animals administered with insulin-loaded nanospheres had a significantly 
reduced serum glucose levels, with respect to the control animals and this 
effect lasted over 6 hrs. 
Cyclosporine A (CyA) is another peptide which has been studied for transport 
to the gastrointestinal tract using polymeric nanoparticles via the oral route. 
CyA is a potent immunosuppressive agent and is widely used for the inhibition 
of graft rejections in the transplant of organs such as heart, liver, skin, lungs, kidney, 
etc. It is also prescribed in autoimmune diseases such as rheumatoid arthiritis 
and Bechet's disease.104-107 Although various formulations of CyA such as Neoral® 
(solution), Sandimmune® (microemulsion) and SangCyA® (amorphous nanoparticles) 
are being marketed, they are faced with the problem of variable bioavailability, 
and the patient has to be monitored for the blood levels of CyA during 
the regimen.108 One of the earlier efforts to improve the bioavailability of CyA 
was done by preparation of pH sensitive nanoparticles using poly (methacrylic 
acid and methacrylate) copolymer (Eudragit®).22 The results were compared with 
Neoral® (a universal standard for CyA oral bioavailability) formulation in rats. 
Nanoparticles exhibited drug entrapment of >90% for different formulations prepared 
from different types of Eudragit® systems. CyA nanoparticles prepared from 
Eudragit® SI 00, an anionic polymer, demonstrated the highest relative bioavailability 
of 132% with respect to Neoral®. Other polymeric nanoparticles also exhibited 
more than 110% relative bioavailability, except for nanoparticles prepared from 
Eudragit® E100 (CyA-ElOO) which is a cationic polymer. In vitro release studies of 
CyA from different nanoparticle preparation illustrated that all nanoparticle preparation 
showed pH-specific release of CyA at pH 7.4, except for CyA-ElOO nanoparticles 
which released the whole payload at pH 2.0. This proves that major CyA from 
CyA-ElOO was released in the stomach upon oral administration accounting for its 
low relative bioavailability with respect to other nanoparticle preparations. 
Polymeric Nanoparticles for Delivery in the Castro-Intestinal Tract 629 
In another study, Wang et a/.,108 examined hydroxypropyl methylcellulose 
phthalate (HPMCP) polymer nanoparticles loaded with CyA for oral delivery. 
HPMCP is a common enteric coating excipient used in the pharmaceutical industry 
for the enteric coating of the tablets. It dissolves specifically at a pH of 7.4 
and releases the contents in the lower intestine. The investigators used two different 
CyA nanoparticle preparations made from different molecular weight of 
the same polymer. Again, a high encapsulation efficiency of over > 95% was 
observed with the nanoparticle preparation, due to hydrophobicity of the drug. 
CyA nanoparticles made of high molecular weight HPMCP exhibited a relative 
bioavailability of over >115%, and the ones made from lower molecular weight 
exhibiting only 82% relative bioavailability against Neoral®. The difference was 
attributed to the pH-independent property of lower molecular weight polymer 
which released entire payload within the stomach itself, thus inactivating the peptide 
drug. The results from the above studies indicate that pH-sensitive nanoparticles 
loaded with CyA can be designed as new carriers for CyA, which exhibit 
a better pharmacokinetic profile compared with the currently marketed CyA 
formulations. 
Nanoparticles made from cationic polymers have been explored as surface coatings 
to improve the oral bioavailability of CyA.106 Male beagle dogs were orally 
administered with CyA nanoparticles coated with chitosan as the poly cationic surface 
modifier. From the results obtained, it was observed that chitosan coated drug 
nanoparticles showed the highest relative bioavailability of 173% with respect to 
Neoral® oral solution. The results were attributed to two properties of the system: 
(i) cationic polymer facilitated the electrostatic interaction with the negatively 
charged mucosa, and (ii) chitosan coated CyA nanoparticles facilitated the opening 
of the tight junctions of the epithelial cells, thus augmenting the paracellular 
transport pathway. 
A series of investigations have been directed towards preparation and evaluation 
of bioavailability and toxicity profile of CyA-loaded polycaprolactone 
nanoparticles.105,109 The nanoparticles, having a diameter of ~100nm were 
prepared by solvent-evaporation procedure and evaluated for biodistribution, 
immunosuppressive activity and nephrotoxicity. Sandimmune® was used as the 
standard for this investigation in rats following oral administration. A significantly 
higher tissue (especially kidney) concentration of CyA was achieved with nanoparticles 
formulations, compared with the solution indicating probability of a higher 
nephrotoxicity. However, further toxicological evaluation with kidney function 
tests indicated no difference in the profiles of two formulations. In vitro lymphocyte 
proliferative activity (an indication of immunosuppressive potential) also showed 
better activity for nanoparticle formulations of comparable doses. The conclusion 
of the investigation was that the nanoparticles formulations can be effective at 
630 Bhavsar, Shenoy & Amiji 
lower dose levels, compared with the solution form and thus may help to reduce 
drug-associated tissue damage. 
Cho et al.m developed several different oral CyA nanoparticle formulations 
consisting of one alkanol solvent and a polyoxyalkylene surfactant, and tested 
them in rats for their bioavailability in comparison to Sandimmune® oral solution. 
Selected formulations based on these pre-clinical investigations were further tested 
for their pharmacokinetic profile in humans. Forty eight healthy males were chosen 
and a randomized, double-blinded, three-way crossover study was conducted with 
Sandimmune® oral solution as standard formulation. From the results obtained, it 
is observed that CyA nanoparticles exhibited a Cmax which was twice as high as 
those achieved by Sandimmune® oral solution and the Tmax was much shorter for 
CyA nanoparticles compared with the standard one. Also, the AUC observed for 
nanoparticle formulations was significantly higher than the standard formulation. 
Polymeric nanoparticles, because of their ability to effectively transport active 
molecules across the gastrointestinal tract have been studied as delivery systems 
for gene therapy and vaccination.111-113 Chen et al™ used DNA-complexed with 
chitosan for transfection of erythropoietin gene to the intestinal epithelium of mice. 
Erythropoietin is a glycoprotein, which stimulates production of red blood cells. 
Erythropoietin is used in patients with anemia associated with chronic renal failure, 
and in cancer patients for simulation of erythropoieisis. Chitosan nanoparticles, 
containing plasmid DNA encoding for erythropoietin (mEpo), were administered 
orally to one group of mice along with other appropriate control dosage forms. 
Erythropoietin gene expression was registered every two days by measuring the 
hematocrit of the mice. Mice which were administered with chitosan loaded mEpo 
showed a 15% increase in hematocrit over other dosage forms, indicating successful 
transfection of mEpo gene across the intestinal epithelium. These results suggests 
that chitosan nanoparticles were able to prevent the mEpo from degradation 
against DNAses and hence the possibility of using them as gene delivery vehicles 
via the oral route. In another study, nanoparticles prepared from cationic biopolymers 
(chitin, chitosan and their derivatives) were proposed to be the carriers for 
oral administration of bioactive compounds for gene therapy.112 The nanoparticles 
with encapsulated plasmid DNA encoding for human coagulation factor IX 
(pFIX) were prepared. The molecular weight of the cationic biopolymers ranged 
from 5 to 200 kDa. The nanoparticles in the size range of 100-200 nm were generated 
by the complex coacervation method and were used for oral administration to 
mice. Human factor IX was detected in the systemic circulation of the mice within 
3 days following oral delivery, but declined after 14 days. The investigators also 
demonstrated the bioactivity of the factor IX transgene product in factor IX knockout 
mice. Heamophilia B is an X-linked bleeding disorder caused by a mutation in 
the factor IX gene. After orally feeding Factor IX transgene-loaded nanoparticles to 
Polymeric Nanoparticles for Delivery in the Castro-Intestinal Tract 631 
the knock-out mice, the clotting time was reduced from 3.5 min to 1.3 min, which 
was comparable with the clotting time of 1 min observed with wild-type mice. 
The investigators proposed that intestinal epithelium was the site of nanoparticle 
absorption and transfection. 
A range of polycationic polymers including gelatin, chitosan, polylysine, polyarginine, 
protamine, speramine, spermidine and polysaccharides could be used 
to prepare the coacervates of the nucleic acids which result in the formation of 
discrete nanoparticles. Roy et al.ni,n5 used such coacervates for effective vaccination 
by the oral route. Chitosan nanoparticles in the size range of 100-200 nm 
were prepared by salting-out technique with the plasmid DNA (pArah2), which 
encodes for the peanut allergen Arah2. The nanoparticles were orally fed into the 
mice and the serum and fecal levels of IgG or IgA were measured periodically. 
High levels of anti-arah2 IgG were observed in the titer of the group which was 
fed with low molecular weight chitosan nanoparticles housing the plasmid DNA 
(pDNA), compared with other groups which were administered with high molecular 
weight chitosan nanoparticles, with or without booster dose. The mice from 
all groups were challenged with crude peanut extracts four weeks after the booster 
dose and positive antibody response were detected in groups immunized by DNA 
nanospheres. These results suggest that chitosan-pDNA nanoparticles delivered 
through the oral route can modify the immune system in mice and protect against 
food allergen induced hypersensitivity. 
Kim et al.u6 prepared PLGA nanoparticles housing H. pylori lysates by solventevaporation 
method. These nanoparticles were administered orally into mice and 
antibody induction was assayed in serum and gastrointestinal tract. Serum IgG 
subclasses were determined by ELISA. The mean antibody titers for serum IgG 
and gut IgA responses were significantly higher than those of the groups immunized 
with the soluble antigen alone. Cholera toxin (CT — a well-established potent 
mucosal adjuvant)-H. pylori had a higher antibody titer compared with PLGA-H. 
pylori nanoparticles. The results of this study indicates that PLGA-H. pylori nanoparticles 
could stimulate H. pylori-specific mucosal and systemic immune responses in 
mice, and also that nanoparticles can be used for vaccination against H. pylori. 
Spray-dried PLGA nanoparticles have been investigated for the oral delivery 
of amifostine.117 Amifostine is an organic thiophosphate prodrug and is dephosphorylated 
by alkaline phosphatase in the tissue to the active free thiol metabolite. 
The major drawback of the drug is that it cannot be administered orally in 
an active form and when administered systemically, it is rapidly cleared from the 
body. PLGA nanoparticles containing amifostine were administered to mice orally 
and tissue distribution was observed for the administered dose. Within 30 min postoral 
administration, the drug was detected in almost all the tissues including blood, 
brain, spleen, kidney, muscle and liver. 
632 Bhavsar, Shenoy & Amiji 
Wheat gram agglutinin (WGA) lectin-functionalized PLGA nanoparticles have 
been successfully prepared and used to encapsulate isoniazid, rifampicin and 
pyrazinamide, which are the three frontline drugs employed in the treatment of 
tuberculosis.118 These PLGA nanoparticles encapsulating antitubecular drugs at 
therapeutic dosage were administered for their in vivo drug disposition studies 
to guinea pigs which were previously infected with Mycobacterium tuberculosis to 
develop the infection. Results obtained for plasma concentration of different drugs 
suggested that PLGA-nanoparticles helped to improve the plasma residence time 
of different drugs after oral/nebulized administration. Rifampicin was detected for 
6 to 7 days in the plasma after oral/aerosolized administration of PLGA-NP, when 
compared with free drug which was detected only for 1 day. Similarly, isoniazid and 
pyrazinamide were maintained for more than 12 days in plasma, compared with a 
single day for the free drug. The presence of these drugs in the tissues such as liver, 
lungs and spleen for a long time favors its application against tuberculosis where 
infection is largely localized in the tissues. Chemotherapeutic studies revealed that 
three doses of oral/aerosolized lectin-coated nanoparticles for 15 days could yield 
undetectable mycobacterial colony forming units, compared with 45 days of oral 
administration of the free drug to achieve the same results. This study suggests 
that polymeric nanoparticles could be favorably used for the effective treatment of 
tuberculosis. 
Popescu etal.19 have proposed the use of biodegradable nanoparticles, prepared 
from naturally occurring polymers such as chitosan, dextran sulfate, dermatan sulfate, 
chondroitin sulfate, keratin sulfate etc. for oral delivery of highly cationic 
active compounds which are highly hydrophilic and could be substrates for PGP. 
Such active compounds include the likes of aminoglycosides, polypeptides, proteins, 
terefenamate, proglumetacin, tiaramide, apazone, etc. Currently, there are no 
technologies for delivery of hydrophilic, cationic drugs by oral administration.19 As 
an example, we will consider streptomycin, which was loaded to chitosan nanoparticles 
and tested for in vivo efficacy using M. tuberculosis infected mice. Streptomycin 
was successfully loaded with an encapsulation efficiency of 50% or higher, with a 
minimal drug loading of 30% w/w of polymer. After oral administration of these 
chitosan nanoparticles in mice a one logio reduction in colony-forming units of the 
bacilli was achieved, compared with the control group. These results show that the 
nanoparticles-based technology can be a break-through for the oral administration 
of aminoglycoside antibiotics, which are otherwise inactive via oral route. 
6.4. Nanoparticles for colon-specific delivery 
The large intestine, which represents the last segment of the gastrointestinal tract 
can suffer from two major inflammatory bowel diseases which are ulcerative colitis 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 633 
and Crohn's disease. Ulcerative colitis occurs more in the distal segment of the large 
intestine and Crohn's disease develops over a very large area of the colon, approximately 
40%. Very little is known about the patho-mechanisms involved in both 
the disease.119-122 Conventionally, treatment of these diseases involves daily intake 
of anti inflammatory drugs which include 5-aminosalicylic acid formulations, glucocorticoids 
and immunosuppressive drugs such as azathioprine, which is taken 
along with methotrexate.123 The major draw back with these conventional formulations 
is that they have to be taken at high doses daily by the oral route, resulting in 
the absorption of these compounds by the small intestine causing possibly strong 
and undesirable effects.17 
Several strategies have been employed for the development of oral delivery 
system for the transport of drugs to the inflamed sites in the colon. These 
include sustained release devices such as prodrugs, macroscopic systems such as 
pH-controlled drug release systems, time-controlled drug release system, enzyme 
controlled drug release systems and also microsized delivery forms such as microspheres 
and nanoparticles. The pH-controlled system relies on the physiological 
difference in the pH of the acidic stomach and that of the distal small intestine, 
time-controlled drug release occurs after a predetermined time lag which is similar 
to the transit time of the system in the small intestine and it ensures delivery of 
the drug into the large intestine. Enzyme-controlled release systems make use of 
the variety of enzymes that are produces by the colonic mucosa to achieve colon 
specific drug delivery. These prodrugs and controlled release devices also have the 
risk of causing adverse side effects, which might result from systemic absorption 
of drug which might occur due to non-specific delivery of the drug all over the 
colon.17'18-123"125 
Polymeric nanoparticles offer an attractive advantage over these systems in that 
they are preferentially absorbed by the mucosal cells of the colon based on their size. 
Mucoadhesion is another property of the polymers, which could be used for site 
specific delivery of the drug to the colon. Polymers such as polysaccharides, which 
include chondroitin sulfate, pectin, dextran and guar gum, have been researched 
for their use as colon specific systems. Chitosan, which is one of the most abundant 
natural polysaccharide, has also been investigated for the development of colon 
specific delivery system due to its well known mucoadhesive properties. A recent 
study by Zhang et al.114 on rats also showed that chitosan gets degraded by the 
cecal and colonic enzymes. Factors which affected the degradation of chitosan in 
colon include its molecular weight and the degree of acetylation. 
A size-dependent bioadhesion of nanoparticles and microparticles in the 
inflamed colonic mucosa has been demonstrated.17 Commercially manufactured 
fluorescent polystyrene particles of different sizes including 100,1000 and 10,000 nm 
were used in the study. The experiments were conducted in rats, which were rectally 
634 Bhavsar, Shenoy & Amiji 
catheterized and treated with trinitrobenzenesulfonic acid (TNBS), for inducing 
inflammatory bowel disease. Polystyrene particles were administered orally to 
the rats and were assessed for localization and deposition of the particles in the 
GIT. Myeloperoxidase (MPO) activity was determined to ensure and quantify the 
inflammation in the colonic area. Size-dependent particle deposition was found in 
the gastrointestinal tract of control group and also in the inflamed tissue. It was 
found that lower size particles exhibited higher incidence of particle deposition 
in the inflamed tissue, with the lowest particle size of 100 nm showing a 6.5-fold 
increase in percentage particle binding, when compared with particle binding of 
the same size in the healthy control group. The overall distribution of the nanoparticles 
in the GIT was assessed by confocal laser scanning microscopy and again it 
was found that 100 nm particles had a higher percentage of localization (38.6%) in 
the mucus of the inflamed tissue, compared with 31.1 % for 1000 nm and only 13.4% 
for 10,000 nm particles. This study proves that nanoparticles are better localized 
and deposited by the macrophages of the inflamed tissue, and that size-dependent 
deposition of particles in the inflamed tissue should be given importance, when 
designing a nanoparticle carrier system for inflammatory bowel disease. 
The same group developed a biocompatible and biodegradable nanoparticle 
system for targeted oral delivery to the inflamed tissues of the colon for patients 
suffering from inflammatory bowel disease using PLGA.18 Two different molecular 
weights of PLGA (5000 and 20,000) were used to prepare nanoparticles containing 
rolipram, an anti inflammatory drug. Emulsification-solvent-evaporation method 
was used for nanoparticles synthesis to yield particle size of less than 500 nm, with 
an encapsulation efficiency of > 80%. Colonic inflammations were induced into the 
rats using TNBS and were checked for the severity of colitis by measuring MPO 
activity. PLGA nanoparticles were orally administered to the rats daily for five days 
and the control group received only saline. PLGA nanoparticles exhibited a local 
anti inflammatory effect by controlled drug release and also proved to be as efficient 
as the free drug in decreasing inflammation of the colitis. Charged interactions of the 
negatively charged PLGA nanoparticles (MW — 20,000) and the positively charged 
proteins of the ulcerated tissue showed a further enhancement of the binding of 
these nanoparticles to the inflamed tissue. 
7. Integrating Polymeric Nanoparticles and Dosage Forms 
If the development of a nanoparticles-based formulation for a drug is a scientifically 
stimulating job, then development of the means to administer them orally to 
humans is a challenging art. As the scientific community is currently busy solving 
the problems associated with the former "scientific" portion, we would like to 
project a few possible scenarios that could be utilized to develop the "art" of oral 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 635 
Hard Gelatin 
capsules 
Enteric coated 
capsules 
Capsules 
Tablets z Osmotic 
tablets Fast 
dissolving 
Floating 
devices 
Soft Gelatin 
formulations 
Selfemulsifying 
Suspensions 
I 
Liquids / dry 
power for oral 
suspension 
Hydrogels 
Films and 
strips 
Fig. 3. Flowchart showing integration of drug-loaded polymeric nanoparticles and 
conventional dosage forms. 
administration of polymeric nanoparticles. Figure 3 provides an overview of conventional 
formulations that could be used to pack the drug-loaded nanoparticles 
for the purpose of oral administration. 
Most of the methods used for manufacturing of the nanoparticles yield drugloaded 
nanoparticles as suspensions (generally aqueous). If the polymers constituting 
the nanoparticles remain stable in aqueous environment for the proposed 
shelf-life period, then they could be directly packaged as oral suspensions, along 
side suitable additives such as flavors, colors, suspending agents and preservatives. 
This would constitute the simplest oral formulation. 
It may be desirable to freeze dry the original drug-loaded nanoparticles suspension 
to limit degradation and also to reduce the levels of organic solvents used 
during their production. In such cases, the drug loaded nanoparticles shall be available 
as free-flowing powder, with or without added stabilizer (i.e. a secondary 
polymer that is used to prevent aggregation during synthesis of nanoparticles). 
636 Bhavsar, Shenoy & Amiji 
This can be mixed with a standard diluent (e.g. lactose) and directly filled into a 
hard gelatin capsule. Another appealing strategy is to formulate as a dry power 
for oral suspension. In this case, the nanoparticles power can be mixed with excipients, 
including suspending agent, sweetening agent (if necessary), flavors, colors 
and preservatives. The contents are to be suspended in water before ingestion. 
Soft-gelatin capsules are popular among conventional oral dosage forms and 
with the availability of novel excipients (from companies such as Gattefosse), their 
application has been extended to meet specialized needs (e.g. sustained release 
or in-situ gelling systems). The drug-loaded nanoparticles can be suspended in a 
suitable medium (usually oil-based) and filled into soft-gelatin capsules. Special 
properties can be imparted to the system by including gel-forming components or 
self-emulsifying components that generate a unique system upon dissolution of the 
gelatin coat in vivo. An in situ formed gel can incorporate nanoparticles to extend 
the dissolution times, or an emulsion may be designed to promote the absorption 
of drugs from the nanoparticles. 
The delivery module can be made in the tablet form as well.126 However, one has 
to evaluate the deformations of the drug-loaded nanoparticles at the compression 
conditions employed. The nanoparticle-tablet can be designed as a floating system 
to increase gastric resident time, or as a bioadhesive system which would increase 
the contact time and hence results in a sustained release, or even a fast-dissolving 
system to expose the nanoparticles quickly to the GIT for further action. Coupling 
osmotic system for delivery of polymeric nanoparticles could be an interesting 
option to offer a multi-step control over drug availability. 
8. Toxicology and Regulatory Aspects 
The ultimate mission of the regulatory body governing the approval of pharmaceutical 
products in the United States (Food and Drug Administration — FDA) is 
not only to protect, but also to provide improvement of public health by assuring 
the safety and efficacy of the products for human and veterinary use. 
The FDA has taken parallel measures along with the advancements in nanotechnology 
to meet novel demands and challenges. There are reasons for the FDA to take 
special steps in promoting the availability of nanotechnology products for public 
use; it is a rapidly growing area of science and is anticipated to lead in the development 
of novel and sophisticated (possibly complex) applications in drug delivery 
systems. As the FDA only regulates to the "claims" made by a sponsor, it may be 
unaware that nanotechnology is being employed to develop that formulation. 
Nanotechnology has been currently evaluated under FDA's Critical Path Initiative 
to keep in pace with the developments in the pharma/biotech industry. Office of 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 637 
Combination Products of the FDA coordinates the regulatory framework for nanotechnology 
products including nanoparticles, and a dedicated FDA Center has 
been proposed for taking the primary responsibility of the review of applications. 
Many of the FDA regulated products are expected to be influenced and revamped 
under nanotechnology, such as drugs, drug/gene/protein delivery systems, vaccines, 
biotechnology products, medical devices and cosmetics. Historically, the FDA 
has approved many products and formulations containing solid particulate matter 
of nano-size range (< 1000 nm). It is also understood that many of the bioactive compounds 
are reduced to nanosize during the process of bioabsorption and there have 
been no severe safety concerns relating to particle size that have been reported earlier. 
To obtain approval for a nanoparticle-based drug delivery system, the industry 
has to address the following issues127: 
8.1. Safety 
The nano-formulations should be evaluated with respect to toxicological 
screening (pharmacology, clinical and histopathological analysis, absorption/ 
disposition/metabolism/excretion (ADME) parameters, genotoxicity, developmental 
toxicity, irritation studies, immunotoxicology and carcinogenicity) projection 
of potential novel and unanticipated reactions and evaluation of excipients 
effects prior to clinical use. An effort should be made to address the following 
questions127: 
• With a reduction in the particle size, there could be a change in size-specific 
effects on the biological activity of the system. Hence, it is important to address 
the issues such as: 
• Will nanoparticles gain access to tissues and cells that normally would be 
bypassed by larger particles? 
• Once nanoparticles enter tissues, how long do they remain intact and how are 
they cleared? 
• If nanoparticles enter cells, what effects do they have on cellular and tissue 
functions? Would there be different effects in the different cell types? 
• What are the differences in the ADME profile of nanoparticles versus larger 
particles? 
• What preclinical screening tests would be useful to identify potential risks 
(in vitro or in vivo)? 
• Can new technologies such as "omics" help identify potential toxicities and how 
can these methodologies complement current testing requirements? 
• Can nanoparticles gain access to the systemic circulation from the route of 
exposure? If nanoparticles enter cells, is there an effect on cellular functions? 
638 Bhavsar, Shenoy & Amiji 
8.2. Quality of material/characterization 
As new toxicological risks that derive from novel materials and delivery systems 
are identified, new tests will be required to ascertain safety and efficacy. Industry 
and academia need to plan and conduct the research to identify potential risks and 
to develop adequate characterization methodologies. 
• What are the forms in which particles are presented to host, tissues, organs, 
organelles and cells? 
• What are the critical physical and chemical properties, including residual solvents, 
processing variables, impurities and excipients? 
• What are the standard tools used for this characterization? 
• What are the validated assays to detect and quantify nanoparticles in tissues, 
medical products, foods and processing equipment? 
• How do physical characteristics impact product quality and performance? 
• How do we determine long and short-term stability of nanomaterials? 
8.3. En vironmen tal considera tions 
• Can nanoparticles be released into the environment following human and 
animal use? 
• What methodologies would identify the nature and quantify the extent of 
nanoparticle release in the environment? 
• What might be the environmental impact on other species (e.g. animals, fish, 
plants, microorganisms)? 
As the materials and the techniques used to manufacture the novel formulations 
may not have prior art to refer to as a standard, there is an additional burden 
on the pharma/biotech industry to carry out a detailed evaluation of the system to 
generate sufficient database for successful industrialization of the product. Some of 
the industrially relevant criteria include understanding the relationship between 
the physico-chemical properties and product performance, effect of process and 
formulation variables on product characteristics, development of analytical tools 
and specifications to regulate product quality, accelerated stability testing as per 
standard protocols to propose a reliable shelf-life, product scale-up to mass production 
and establishment of manufacturing standards and development of reference 
materials/standards as guidelines for quality assurance. Development of validated 
testing methods/protocols and establishment of reference standards through a thorough 
and logical process remains to be the major responsibility of the industry for 
convincing the FDA to get product approval. 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 639 
While considering the application of a polymeric nanoparticles-based 
formulation, the FDA may want the industry to include evidence for the parameters 
listed below: 
• Particle size and size distribution 
• Surface area, surface chemistry, surface coating and porosity 
• Hydrophilicity and surface charge density 
• Purity and quality 
• Stability (on shelf and upon administration) 
• Manufacturing and Controls 
• Drug release parameters and bioequivalence testing considerations 
9. Conclusion and Outlook 
Under the light of current literature (i.e. articles, books, patents and information 
posted on the nanotech company websites) and the product pipelines of 
leading pharma/biotech companies, it is evident that we would be seeing many 
nanotechnology-based pharmaceutical products in this century. Table 2 lists few 
of the important products in the drug delivery pipeline that are based on polymeric 
nanoparticles. It is likely that the oral formulations would dominate this 
specialized segment of novel dosage forms. The chemical/polymer industry has 
been feeding the drug delivery scientists with a variety of biopolymers, having 
wide range of specialized properties. Nanoparticles made from the biopolymers 
are likely to dominate the novel drug delivery systems in the oral market because of 
the cost-to-benefit ratio, excellent stability, flexibility for industrial production and 
a voluminous database available, with respect to the regulatory issues addressed 
earlier. Polymeric nanoparticles are also being explored for topical applications 
and as sterile dosage forms for ophthalmic, nasal, subcutaneous and intravenous 
applications. 
There are several other potential nanoparticles technologies which fall outside 
the coverage of this chapter, which are based on nanoparticles made from the drugs 
themselves. They are termed as nanosuspensions, nanocrystals or insoluble drug 
delivery technologies.28,128,129 Essentially, all of them are colloidal dispersions of 
pure drug particles that are stabilized by polymers, surfactants or lipids. They are 
synthesized either by physical (e.g. size reduction by milling) or chemical (e.g. 
change in solubility induced by pH or solvent exchange) means in the presence of 
stabilizing agents. The striking advantage of these technologies is the high drug 
loading efficiency and the simplicity associated with its production. These have 
been the first to roll out from the research and development scale to the industrial 
production scale under nanoparticle category (Rapamune® oral solution and tablets 
640 Bhavsar, Shenoy & Amiji 
Table 2 Product pipeline of polymeric nanoparticles (Source: PharmaProjects). 
Company Technology Bioactive Route of delivery 
compound 
Novavax, USA Micellar nanoparticles Testosterone Subcutaneous 
Flamel Technologies, Medusa® nanoparticles Insulin/Interferon Subcutaneous 
France 
BioAUiance, France 
Munich Biotech, 
Germany 
BioSante, USA 
Targesome, USA 
American 
Bioscience, USA 
Advectus Life 
Sciences, Canada 
Nanocarrier, Japan 
Wyeth 
of amino acids 
Polydsohexyl 
cyanoacrylate) 
nanoparticles 
Drug nanoparticles 
Calcium phospahte 
nanoparticles 
Self-assembling lipid 
nanospheres 
Albumin-Drug 
nanoparticles 
Poly(butylcyanoacrylate) 
nanoparticles 
Micellar nanoparticles 
Drug nanoparticles 
Doxorubicin 
Paclitaxel 
Insulin 
Therapeutic/ 
Diagnostic 
Paclitaxel 
Doxorubicin 
Water insoluble 
drugs 
Rapamycin 
Intravenous 
Intravenous 
Oral 
Intravenous 
Intravenous 
Intravenous 
N/A 
Oral 
Pharmaceuticals, 
USA 
containing sirolimus from Wyeth and SangCya® oral solution from SangStat Corporation 
containing CyA). If the science of pharmaceutical product development is 
undergoing a transformation from a traditional pharmaceutics to a more innovative 
molecular or nano-pharmaceutics, the major credit would be taken by a combination 
of polymer based systems and nanoparticles. It is more of a belief than a hope 
that the polymeric nanoparticles would address many of the therapeutic issues that 
are posing hurdles to a formulation scientist in this century. 
References 
1. Wilson K and Waugh A (1996) Anatomy and Physiology in Health and Illness. Churchill 
Livingstone: New York. 
2. Tortora G and Anagnostakos N (1978) Principles of Anatomy and Physiology. Harper and 
Row: New York. 
3. Tate P, Seeley R and Stephens T (1994) Understanding the Human Body. Mosby: St. Louis. 
4. Solomon E (1992) Introduction to Human Anatomy and Physiology. W. B. Saunders 
Company, Philadelphia. 
5. McClintic JR (1980) Basic Anatomy and Physiology of the Human Body. John Wiley and 
Sons: New York. 
Polymeric Nanoparticles for Delivery in the Castro-Intestinal Tract 641 
6. Ganong W (2003) Review of Medical Physiology. Appleton and Lange: Norwalk, CT. 
7. Shojaei A (1998) Buccal mucosa as a route for systemic drug delivery: A review. / Pharm 
Pharma Sci 1:15-30. 
8. Zhang H, Zhang J and Streisand JB (2002) Oral mucosal drug delivery: Clinical pharmacokinetics 
and therapeutic applications. Clin Pharmacokinet 41:661-680. 
9. Collins L and Dawes C (1987) The surface area of the adult human mouth and thichness 
of the salivary film covering the teeth and oral mucosa. / Dent Res 66:1300-1302. 
10. Harris D and Robinson J (1992) Drug delivery via the mucous membranes of the oral 
cavity. / Pharm Sci 81:1-10. 
11. Sakuma S, Hayashi M and Akashi M (2001) Design of nanoparticles composed of graft 
copolymers for oral peptide delivery. Adv Drug Del Rev 47:21-37. 
12. Gandhi RR, J (1994) Oral cavity as a site for bioadhesive drug delivery. Adv Drug Del 
Rev 13:43-74. 
13. Florence AT (1993) Particulate delivery: The challenge of the oral route. Drugs Pharm 
Sci 61:65-107. 
14. Jung T, Kamm W, Breitenbach A, Kaiserling E, Xiao JX and Kissel T (2000) Biodegradable 
nanoparticles for oral delivery of peptides: Is there a role for polymers to affect mucosal 
uptake? Eur } Pharm Biopharm 50:147-160. 
15. Friend DR (2004) Drug delivery to the small intestine. Curr Gastroenterol Rep 6:371-376. 
16. Florence AT, Hussain N and Jani P (1995) Nanoparticles as carriers for oral peptide 
absorption: Studies on particle uptake and fate. / Control Rel 36:39-46. 
17. Lamprecht A, Schafer U and Lehr CM (2001) Size-dependent bioadhesion of microand 
nanoparticulate carriers to the inflamed colonic mucosa. Pharm Res 18:788-793. 
18. Lamprecht A, Ubrich N, Yamamoto H, Schafer U, Takeuchi H,MaincentP, Kawashima Y 
and Lehr CM (2001) Biodegradable nanoparticles for targeted drug delivery in treatment 
of inflammatory bowel disease. / Pharmacol Exp Ther 299:775-781. 
19. Popescu C and Onyuksel H (2004) Biodegradable nanoparticles incorporating highly 
hydrophilic positively charged drugs US Patent No. 20040247683. 
20. Tobio M, Sanchez A, Vila A, Soriano II, Evora C, Vila-Jato JL and Alonso MJ (2000) The 
role of PEG on the stability in digestive fluids and in vivo fate of PEG-PLA nanoparticles 
following oral administration. Coll Surf B Biointerf 18:315-323. 
21. Florence AT (2004) Issues in oral nanoparticle drug carrier uptake and targeting. 
/ Drug Targ 12:65-70. 
22. Dai J, Nagai T, Wang X, Zhang T, Meng M and Zhang Q (2004) pH-sensitive nanoparticles 
for improving the oral bioavailability of cyclosporine A. Int} Pharm 280:229-240. 
23. Takeuchi H, Yamamoto H and Kawashima Y (2001) Mucoadhesive nanoparticulate 
systems for peptide drug delivery. Adv Drug Del Rev 47:39-54. 
24. Vemuri S and Rhodes CT (1995) Preparation and characterization of liposomes as 
therapeutic delivery systems: A review. Pharm Acta Helv 70:95-111. 
25. Uchegbu IF and Vyas SP (1998) Non-ionic surfactant based vesicles (niosomes) in drug 
delivery. Int} Pharm 172:33-70. 
26. Douglas SJ, Davis SS and Ilium L (1987) Nanoparticles in drug delivery. Crit Rev Ther 
Drug Can Syst 3:233-261. 
642 Bhavsar, Shenoy & Amiji 
27. Boas U and Heegaard PM (2004) Dendrimers in drug research. Chem Soc Rev 33:43-63. 
28. Rabinow BE (2004) Nanosuspensions in drug delivery. Nat Rev Drug Discov 3:785-796. 
29. Sonneville-Aubrun O, Simonnet JT and L'Alloret F (2004) Nanoemulsions: A new vehicle 
for skincare products. Adv Coll Interf Sci 108-109:145-149. 
30. Tenjarla S (1999) Microemulsions: An overview and pharmaceutical applications. Crit 
Rev Ther Drug Can Syst 16:461-521. 
31. Kataoka K, Harada A and Nagasaki Y (2001) Block copolymer micelles for drug delivery: 
Design, characterization and biological significance. Adv Drug Del Rev 47:113-131. 
32. Solaro R (2002) Nanostructured polymeric systems in targeted release of proteic drugs 
and in tissue engineering China-EU Forum on Nanosized Technology. China-EU Forum, 
Beijng. 
33. Quintanar-Guerrero DA, E; Fessi, H; Doelker, E (1998) Preparation techniques and 
mechanisms of formation of biodegradable nanoparticles from preformed polymers. 
Drug Dev Ind Pharm 24:1113-1128. 
34. Barratt G (2003) Colloidal drug carriers: Achievements and perspectives. Cell Mol Life 
Sci 60:21-37. 
35. Soppimath KS, Aminabhavi TM, Kulkarni AR and Rudzinski WE (2001) Biodegradable 
polymeric nanoparticles as drug delivery devices. / Control Rel 70:1-20. 
36. Kreuter J (2001) Nanoparticulate systems for brain delivery of drugs. Adv Drug Del Rev 
47:65-81. 
37. Alonso MJ (2001) Polymeric nanoparticles: New systems for improving ocular bioavailability 
of drugs. Arch Soc Esp Oftalmol 76:453-454. 
38. Barratt G (2003) Colloidal drug carriers: Achievements and perspectives. Cell Mol Life 
Sci 60:21-37. 
39. Kreuter J (1996) Nanoparticles and microparticles for drug and vaccine delivery. / Anat 
189:503-505. 
40. Speiser PP (1991) Nanoparticles and liposomes: A state of the art. Meth Find Exp Clin 
Pharmacol 13:337-342. 
41. Panyam J and Labhasetwar V (2003) Biodegradable nanoparticles for drug and gene 
delivery to cells and tissue. Adv Drug Del Rev 55:329-347. 
42. Otsuka H, Nagasaki Y and Kataoka K (2003) PEGylated nanoparticles for biological 
and pharmaceutical applications. Adv Drug Del Rev 55:403^19. 
43. Emerich DF and Thanos CG (2003) Nanotechnology and medicine. Exp Opin Biol Ther 
3:655-663. 
44. Lockman PR, Mumper RJ, Khan MA and Allen DD (2002) Nanoparticle technology for 
drug delivery across the blood-brain barrier. Drug Dev Ind Pharm 28:1-13. 
45. Brigger I, Dubernet C and Couvreur P (2002) Nanoparticles in cancer therapy and 
diagnosis. Adv Drug Del Rev 54:631-651. 
46. Moghimi SM, Hunter AC and Murray JC (2001) Long-circulating and target-specific 
nanoparticles: Theory to practice. Pharmacol Rev 53:283-318. 
47. Lambert G, Fattal E and Couvreur P (2001) Nanoparticulate systems for the delivery 
of antisense oligonucleotides. Adv Drug Del. Rev 47:99-112. 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 643 
48. Janes KA, Calvo P and Alonso MJ (2001) Polysaccharide colloidal particles as delivery 
systems for macromolecules. Adv Drug Del Rev 47:83-97. 
49. Benns JM and Kim SW (2000) Tailoring new gene delivery designs for specific targets. 
/ Drug Targ 8:1-12. 
50. Tan Y, Whitmore M, Li S, Frederik P and Huang L (2002) LPD nanoparticles-novel 
nonviral vector for efficient gene delivery. Meth Mol Med 69:73-81. 
51. Vijayanathan V, Thomas T and Thomas TJ (2002) DNA nanoparticles and development 
of DNA delivery vehicles for gene therapy. Biochemistry 41:14085-14094. 
52. Cui Z and Mumper RJ (2003) Microparticles and nanoparticles as delivery systems for 
DNA vaccines. Crit Rev Ther Drug Carr Syst 20:103-137. 
53. Vila A, Sanchez A, Evora C, Soriano I, Vila Jato JL and Alonso MJ (2004) PEG-PLA 
nanoparticles as carriers for nasal vaccine delivery. / Aerosol Med 17:174-185. 
54. Prego C, Garcia M, Torres D and Alonso MJ (2005) Transmucosal macromolecular drug 
delivery. / Control Rel 101:151-162. 
55. Videira MA, Botelho MF, Santos AC, Gouveia LF, de Lima JJ and Almeida AJ (2002) Lymphatic 
uptake of pulmonary delivered radiolabelled solid lipid nanoparticles. / Drug 
Targ 10:607-613. 
56. Shim J, Seok Kang H, Park WS, Han SH, Kim J and Chang IS (2004) Transdermal 
delivery of mixnoxidil with block copolymer nanoparticles. / Control Rel 97:477-484. 
57. General S and Thunemann AF (2001) pH-sensitive nanoparticles of poly (amino acid) 
dodecanoate complexes. Int} Pharm 230:11-24. 
58. Mansouri S, Lavigne P, Corsi K, Benderdour M, Beaumont E and Fernandes JC 
(2004) Chitosan-DNA nanoparticles as non-viral vectors in gene therapy: Strategies 
to improve transfection efficacy. Eur } Pharm Biopharm 57:1-8. 
59. Thunemann AF and General S (2001) Nanoparticles of a polyelectrolyte-fatty acid 
complex: Carriers for Q10 and triiodothyronine. / Control Rel 75:237-247. 
60. General S, Rudloff J and Thunemann AF (2002) Hollow nanoparticles via stepwise 
complexation and selective decomplexation of poly (ethylene imine). Chem Commun 
5:534-535. 
61. Akiyoshi K, Kobayashi S, Shichibe S, Mix D, Baudys M, Kim SW and Sunamoto J (1998) 
Self-assembled hydrogel nanoparticle of cholesterol-bearing pullulan as a carrier of 
protein drugs: Complexation and stabilization of insulin. / Control Rel 54:313-320. 
62. Sang Yoo H and Gwan Park T (2004) Biodegradable nanoparticles containing proteinfatty 
acid complexes for oral delivery of salmon calcitonin. } Pharm Sci 93:488-495. 
63. Du J, Zhang S, Sun R, Zhang LF, Xiong CD and Peng YX (2005) Novel polyelectrolyte carboxymethyl 
konjac glucomannan-chitosan nanoparticles for drug delivery. II. Release 
of albumin in vitro. J Biomed Mater Res B Appl Biomater 72:299-304. 
64. Sun YP, Meziani MJ, Pathak P and Qu L (2005) Polymeric nanoparticles from rapid 
expansion of supercritical fluid solution. Chemistry 11:1366-1373. 
65. Young TJ, Johnson KP, Pace GW and Mishra AK (2004) Phospholipid-stabilized 
nanoparticles of cyclosporine A by rapid expansion from supercritical to aqueous solution. 
AAPS PharmSciTech 5:E11. 
644 Bhavsar, Shenoy & Amiji 
66. Subramaniam B, Rajewski RA and Snavely K (1997) Pharmaceutical processing with 
supercritical carbon dioxide. / Pharm Sci 86:885-890. 
67. Gokhale AA, Khushi B, Dave RN and Pfeffer R (2005) Formation of polymer nanoparticles 
in supercritical fluid jets. Nanotech (May). 
68. Couvreur P (1988) Polyalkylcyanoacrylates as colloidal drug carriers. Crit Rev Ther 
Drug Carr Syst 5:1-20. 
69. Brannon-Peppas L (1997) Polymers in controlled drug delivery. Med Plastics Biomater 
Mag, 34. 
70. Kumar MNV (2000) Nano and microparticles as controlled drug delivery devices 
/ Pharm Pharmaceut Sci 3:234-258. 
71. McCarron PA, Donnelly RF, Canning PE, McGovern JG and Jones DS (2004) Bioadhesive, 
non-drug-loaded nanoparticles as modulators of candidal adherence to buccal 
epithelial cells: A potentially novel prophylaxis for candidosis. Biomaterials 25:2399- 
2407. 
72. Ch'ng HS, Park H, Kelly P and Robinson JR (1985) Bioadhesive polymers as platforms 
for oral controlled drug delivery II: Synthesis and evaluation of some swelling, waterinsoluble 
bioadhesive polymers. / Pharm Sci 74:399-405. 
73. Leung SR, J. (1990) Polymer structure features contributing to mucoadhesion: II. 
J Control Rel 12:187-194. 
74. Sanzgiri Y, Topp E, Benedetti L and Stella V (1994) Evaluation of mucoadhesive properties 
of hyaluronic acid benzyl esters. Int J Pharm 107:91-97. 
75. Luessen HL, Verhoef JC, Borchard G, Lehr CM, de Boer AG and Junginger HE (1995) 
Mucoadhesive polymers in peroral peptide drug delivery. II. Carbomer and polycarbophil 
are potent inhibitors of the intestinal proteolytic enzyme trypsin. Pharm Res 
12:1293-1298. 
76. Park KR, J. (1984) Bioadhesive polymers as platforms for oral-controlled drug delivery: 
Method to study bioadhesion. Int J Pharm 19:107-127. 
77. Nagai TM, Y (1993) Buccal delivery systems using hydrogels. Adv Drug Del Rev 
11:179-191. 
78. Taylan B, Yilmaz C, Guven O, Kes S and Hincal A (1996) Design evaluation of sustainedrelease 
and buccal adhesive propranolol hydrochloride tablets. / Control Rel 38:11-20. 
79. Leung S and Robinson J (1988) The contribution of anionic polymer structural features 
to mucoadhesion. / Control Rel 5:223-231. 
80. Venugopalan P, Sapre A, Venkatesan N and Vyas SP (2001) Pelleted bioadhesive polymeric 
nanoparticles for buccal delivery of insulin: Preparation and characterization. 
Pharmazie 56:217-219. 
81. Gandhi R and Robinson J (1988) Bioadhesion in drug delivery. Indian J Pharm Sci 
50:145-152. 
82. Neuberger W (2004) Treatment of periodontal disease with photosensitizers, Int Patent 
WO 2004/024080 A2. 
83. Chowdhary RK and Dolphin D (2002) Supports for photosensitizer formulations, US 
Patent No. 20020061330. 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 645 
84. Hejazi R and Amiji M (2002) Stomach-specific anti-H. pylori therapy. I: Preparation 
and characterization of tetracyline-loaded chitosan microspheres. Int } Pharm 
235:87-94. 
85. Umamaheshwari RB, Ramteke S and Jain NK (2004) Anti-Helicobacter pylori effect of 
mucoadhesive nanoparticles bearing amoxicillin in experimental gerbils model. AAPS 
PharmSciTech 5:1-9. 
86. Ezpeleta I, Arangoa MA, Irache JM, Stainmesse S, Chabenat C, Popineau Y and 
Orecchioni AM (1999) Preparation of Ulex europaeus lectin-gliadin nanoparticle conjugates 
and their interaction with gastrointestinal mucus. Int} Pharm 191:25-32. 
87. Dawson M, Krauland E, Wirtz D and Hanes J (2004) Transport of polymeric nanoparticle 
gene carriers in gastric mucus. Biotechnol Prog 20:851-857. 
88. Gao H, Wang JY, Shen XZ, Deng YH and Zhang W (2004) Preparation of magnetic 
polybutylcyanoacrylate nanospheres encapsulated with aclacinomycin A and its effect 
on gastric tumor. World } Gastroenterol 10:2010-2013. 
89. Ponchel G and Irache J (1998) Specific and non-specific bioadhesive particulate 
systems for oral delivery to the gastrointestinal tract. Adv Drug Del Rev 34: 
191-219. 
90. Jani P, Halbert GW, Langridge J and Florence AT (1990) Nanoparticle uptake by the rat 
gastrointestinal mucosa: Quantitation and particle size dependency. ] Pharm Pharmacol 
42:821-826. 
91. Araujo L, Sheppard M, Lobenber Rand Kreuter J (1999) Uptake of PMMAnanoparticles 
from the gastrointestinal tract after oral administration to rats: Modification of the body 
distribution after suspension in surfactant solutions and in oil vehicles. Int J Pharm 
176:209-224. 
92. Couvreur P, Lenaerts V, Kante B, Roland M and Speiser P (1980) Oral and parentral 
administration of insulin associated hydrolysable nanoparticles. Acta Pharm Technol 
26:220-222. 
93. Delie F and Blanco-Prieto M (2005) Polymeric particulates to improve oral bioavailability 
of peptide drugs. Molecules 10:65-80. 
94. Narayani R (2001) Oral delivery of insulin — making needles needless. Trends Biomater 
Artif Organs 15:12-16. 
95. Sakuma S, Suzuki N, Sudo R, Hiwatari K, Kishida A and Akashi M (2002) Optimized 
chemical structure of nanoparticles as carriers for oral delivery of salmon calcitonin. 
Int ] Pharm 239:185-195. 
96. Damge C, Michel C, Aprahamian M and Couvreur P (1988) New approach for oral 
administration of insulin with polyalkylcyanoacrylate nanocapsules as drug carrier. 
Diabetes 37:246-251. 
97. Pan Y, Li YJ, Zhao HY, Zheng JM, Xu H, Wei G, Hao JS and Cui FD (2002) Bioadhesive 
polysaccharide in protein delivery system: Chitosan nanoparticles improve the 
intestinal absorption of insulin in vivo. Int} Pharm 249:139-147. 
98. Damge C, Michel C, Aprahamian M, Couvreur P, Devissaguet J (1990) Nanocapsules 
as carriers for oral peptide delivery. / Control Rel 13:233-239. 
646 Bhavsar, Shenoy & Amiji 
99. Aboubakar M, Couvreur P, Pinto-Alphandary H, Gouritin B, Lacour B, Farinotti R, 
Puisieux F and Vauthier C (2000) Insulin-loaded nanocapsules for oral administration: 
In vitro and in vivo investigation. Drug Dev Res 49:109-117. 
100. Ramtoola Z (1997) Controlled release biodegradable nanoparticles containing insulin. 
US Patent No. 5641515. 
101. Agnihotri SA, Mallikarjuna NN and Aminabhavi TM (2004) Recent advances on 
chitosan-based micro- and nanoparticles in drug delivery. / Control Rel 100:5-28. 
102. Carino GP, Jacob JS and Mathiowitz E (2000) Nanosphere based oral insulin delivery. 
/ Control Rel 65:261-269. 
103. Foss AC, Goto T, Morishita M and Peppas NA (2004) Development of acrylic-based 
copolymers for oral insulin delivery. Eur } Pharm Biopharm 57:163-169. 
104. Lee WK, Park JY, Yang EH, Suh H, Kim SH, Chung DS, Choi K, Yang CW and 
Park JS (2002) Investigation of the factors influencing the release rates of cyclosporin 
A-loaded micro- and nanoparticles prepared by high-pressure homogenizer. / Control 
Rel 84:115-123. 
105. Molpeceres L, Aberturas MR and Guzman M (2000) Biodegradable nanoparticles as 
a delivery system for cyclosporine: Preparation and characterization. / Microencapsul 
17:599-614. 
106. El-Shabouri MH (2002) Positively charged nanoparticles for improving the oral 
bioavailability of cyclosporin-A. Int} Pharm 249:101-108. 
107. Young TJ, Johnston KP, Pace GW and Mishra AK (2003) Phospholipid-stablized 
nanoparticles of cyclosporine A by rapid expansion from supercritical to aqueous 
solution. AAPS PharmSciTech 5:1-16. 
108. Wang XQ, Dai JD, Chen Z, Zhang T, Xia GM, Nagai T and Zhang Q (2004) Bioavailability 
and pharmacokinetics of cyclosporine A-loaded pH-sensitive nanoparticles for oral 
administration. / Control Rel 97:421-429. 
109. Varela MC, Guzman M, Molpeceres J, del Rosario Aberturas M, Rodriguez-Puyol D 
and Rodriguez-Puyol M (2001) Cyclosporine-loaded polycaprolactone nanoparticles: 
Immunosuppression and nephrotoxicity in rats. Eur } Pharm Sci 12:471^178. 
110. Cho M, Levy R, Pouletty P, Floc'h R and Merle C (1997) Oral cyclosproin formulations, 
Int Patent WO 97/07787. 
111. Chen J, Yang WL, Li G, Qian J, Xue JL, Fu SK and Lu DR (2004) Transfection of mEpo gene 
to intestinal epithelium in vivo mediated by oral delivery of chitosan-DNA nanoparticles. 
World } Gastroenterol 10:112-116. 
112. Leong K, Okoli G and Hottelano G (2003) Compositions for oral gene therapy and 
methods of using same. International Patent WO 03/02867 A2. 
113. Chang SF, Chang HY, Tong YC, Chen SH, Hsaio FC, Lu SC and Liaw J (2004) Nonionic 
polymeric micelles for oral gene delivery in vivo. Hum Gene Ther 15:481-493. 
114. Roy K, Shau-Ku H, Sampsom H and Leong K (2002) Oral delivery of nucleic acid 
vaccines by particulate complexes. US Patent 6475995 Bl. 
115. Roy K, Mao HQ, Huang SK and Leong KW (1999) Oral gene delivery with chitosan- 
DNA nanoparticles generates immunologic protection in a murine model of peanut 
allergy. Nat Med 5:387-391. 
Polymeric Nanoparticles for Delivery in the Gastro-lntestinal Tract 647 
116. Kim SY, Doh HJ, Jang MH, Ha YJ, Chung SI and Park HJ (1999) Oral immunization with 
Helicobacter pylori-loaded poly(D, L-lactide-co-glycolide) nanoparticles. Helicobacter 
4:33-39. 
117. Pamujula S, Graves RA, Freeman T, Srinivasan V, Bostanian LA, Kishore V and T.K. M 
(2004) Oral delivery of spray dried PLGA/amifostine nanoparticles. J Pharm Pharmacol 
56:1119-1125. 
118. Sharma A, Sharma S and Khuller GK (2004) Lectin-functionalized poly (lactide-coglycolide) 
nanoparticles as oral/aerosolized antitubercular drug carriers for treatment 
of tuberculosis. / Antimicrob Chemother 54:761-766. 
119. Allison MC, Cornwall S, Poulter LW, Dhillon AP and Pounder RE (1988) Macrophage 
heterogeneity in normal colonic mucosa and in inflammatory bowel disease. Gut 
29:1531-1538. 
120. Seldenrijk CA, Drexhage HA and Meuwissen SGM (1989) Dendritic cells and scavanger 
macrophage in chronic inflammatory bowel disease. Gut 30:484-491. 
121. Probert CS, Chott A, Turner JR, Saubermann LJ, Stevens AC, Bodinaku K, Elson CO, 
Balk SP and Blumberg RS (1996) Persistent clonal expansions of peripheral blood CD4+ 
lymphocytes in chronic inflammatory bowel disease. / Immunol 157:3183-3191. 
122. Tabata Y, Inoue Y and Ikada Y (1996) Size effect on systemic and mucosal immune 
responses induced by oral administration of biodegradable microspheres. Vaccine 
14:1677-1685. 
123. Lamprecht A, Stallmach A, Kawashima Y and Lehr CM (2002) Carrier systems for the 
treatment of inflammatory bowel disease. Drugs Put 27:961-971. 
124. Zhang H, Alsarra IA and Neau SH (2002) An in vitro evaluation of a chitosan-containing 
multiparticulate system for macromolecule delivery to the colon. Int J Pharm 239: 
197-205. 
125. Rodriguez M, Vila-Jato JL and Torres D (1998) Design of a new multiparticulate system 
for potential site-specific and controlled drug delivery to the colonic region. / Control 

Rel 55:67-77. 
126. Schmidt C and Bodmeier R (1999) Incorporation of polymeric nanoparticles into solid 
dosage forms. / Control Rel 57:115-125. 
127. Refer to presentations and publications at www.fda.gov/nanotechnology. 
128. Refer to the technical information posted by Skye Pharma at www.skyepharma.com/ 
solubilization.html. 
129. Refer to the technical information posted by Elan at http://www.elan.com/drugdelivery/
drug delivery/nanocrystal_technology.asp. 
130. Arangoa MA, Campanero MA, Renedo MJ, Ponchel G and Irache JM (2001) Gliadin 
nanoparticles as carriers for the oral administration of lipophilic drugs. Relationships 
between bioadhesion and pharmacokinetics. Pharm Res 18:1521-1527. 
131. Murakami HK, M.; Takeuchi, H.; Kawashima, Y. (2000) Utilization of poly(DL-lactideco-
glycolide) nanaoparticles for preparation of mini depot tablets by direct compression. 
/ Control Rel 37:29-36. 
132. Tobio M, Sanchez A, Vila A, Soriano II, Evora C, Vila-Jato JL and Alonso MJ 
(2000) The role of PEG on the stability in digestive fluids and in vivo fate of 
648 Bhavsar, Shenoy & Amiji 
PEG-PLA nanoparticles following oral administration. Coll Surf B Biointerf 18: 
315-323. 
133. Hillery A, Florence A (1996) The effect of adsorbed poloxamer 188 and 407 surfactants 
on the intestinal uptake of 60-nm polystyrene particles after oral administration in the 
rat. Int} Pharm 132:123-130. 
134. Sakuma S, Sudo R, Suzuki N, Kikuchi H, Akashi M, Ishida Y and Hayashi M (2002) 
Behavior of mucoadhesive nanoparticles having hydrophilic polymeric chains in the 
intestine. / Control Rel 81:281-290. 
135. Yoo HP, T. (2004) Biodegradable nanoparticles containing protein-fatty acid complexes 
for oral delivery of salmon calcitonin. / Pharm Sci 93:488^95. 
136. Arbos P, Campanero MA, Arangoa MA and Irache JM (2004) Nanoparticles with specific 
bioadhesive properties to circumvent the pre-systemic degradation of fluorinated 
pyrimidines. / Control Rel 96:55-65. 
137. Jiao Y, Ubrich N, Marchand-Arvier M, Vigneron C, Hoffman M, Lecompte T and 
Maincent P (2002) In vitro and in vivo evaluation of oral heparin-loaded polymeric 
nanoparticles in rabbits. Circulation 105:230-235. 
138. Arbos P, Campanero MA, Arangoa MA, Renedo MJ and Irache JM (2003) Influence of 
the surface characteristics of PVM/MA nanoparticles on their bioadhesive properties. 
/ Control Rel 89:19-30. 
139. Barichello JM, Morishita M, Takayama K and Nagai T (1999) Encapsulation of 
hydrophilic and lipophilic drugs in PLGA nanoparticles by the nanoprecipitation 
method. Drug Dev Ind Pharm 25:471-476. 
140. Molpeceres J, Aberturas MR and Guzman M (2000) Biodegradable nanoparticles as 
a delivery system for cyclosporine: Preparation and characterization. / Microencapsul 
17:599-614. 
141. Ubrich N, Schmidt C, Bodmeier R, Hoffman M and Maincent P (2005) Oral evaluation in 
rabbits of cyclosporin-loaded Eudragit RS or RL nanoparticles. Int}Pharm 288:169-175. 
142. Mesiha MS, Sidhom MB and Fasipe B (2005) Oral and subcutaneous absorption of 
insulin poly(isobutylcyanoacrylate) nanoparticles. Int ] Pharm 288:289-293. 
143. Pamujula S, Graves RA, Freeman T, Srinivasan V, Bostanian LA, Kishore V and 
Mandal TK (2004) Oral delivery of spray dried PLGA/amifostine nanoparticles. 
} Pharm Pharmacol 56:1119-1125. 
27 
Nanoparticular Carriers for Ocular 
Drug Delivery 
Alejandro Sanchez and Maria J. Alonso 
The major goal in ocular drug delivery is to obtain therapeutic drug concentrations 
at the intended site of action (i.e. located at the eye surface or in the inner 
eye), for reasonable periods of time. The strategies explored towards this goal have 
been (i) the design of topical ocular delivery systems which promote the concentration 
of the drug on the eye surface, and, if necessary, facilitate the drug transfer 
from the extraocular tissues to the internal structures of the eye; (ii) the design of 
injectable controlled release systems which deliver the drug directly to the sclera 
(subconjunctival injection) or to the internal structures of the eye (intravitreal injection), 
for extended periods of time. Among the delivery systems designed so far 
for these purposes, those of a nanoscale size are particularly attractive from the 
point of view of easiness of administration and patient acceptability, since they 
can be applied in the form of a non-viscous liquid. This chapter aims to describe 
the advances and the actual potential of polymer-based nanostructures such as 
nanoparticles and nanocapsules, for topical ocular drug delivery. Since the complexity 
of these nanostructures has increased over the time, these nanostructures 
have been classified into first, second and third-generation nanocarriers. Additionally, 
the last sections of the chapter were intended to present the possibility to use 
nanoparticulate drug carriers for injection (i.e. subconjunctival, intravitreal), and to 
underline the specific advantages of nanosystems over large dimensional devices 
for intraocular drug delivery. Overall, this review chapter shows the great potential 
649 
650 Sanchez & Alonso 
that nanosystems offer in terms of improving the efficacy of drugs used in ocular 
therapies. Moreover, it emphasizes that the advances achieved in the understanding 
of the interaction of nanosystems with the ocular tissues should, logically, result 
in the design of sophisticated systems specifically tailored for ocular drug delivery. 
1. Biopharmaceutical Barriers in Ocular Drug 
Delivery. Classification of Nanoparticulate 
Carriers for Ocular Drug Delivery 
Unlike other routes described in previous chapters of this book, the different modalities 
of ocular administration (i.e. topical, subconjuctival and intravitreal) are exclusively 
intended to deliver drugs locally for the treatment of ophthalmic processes, 
and not as an entry to the systemic circulation. Among these modalities, the topical 
ocular administration is the easiest and best accepted by the patients. Liquid formulations, 
solutions and suspensions, are the most commonly applied for topical 
ocular administration, since they are easy to use and do not interfere with vision. 
However, these formulations are often quite ineffective due to the defense mechanisms 
of the ocular apparatus. These mechanisms have been described in detail in 
several review articles and text-books.1-3 Firstly, most of the drug applied topically 
onto the eye is immediately diluted in the precorneal tear film. The excess fluid 
spills over the lid margin and the remainder is rapidly drained into the nasolachrymal 
duct. As a consequence, most of the applied drug solution is cleared within 
2-4 min.4'5 In addition, a proportion of the drug will not be available for therapeutic 
action at the ocular level, but will be absorbed to the systemic circulation through 
the surrounding extraorbital tissues, mainly the conjunctiva (unproductive drug 
absorption). On the other hand, in the case of drugs whose target is located in the 
inner eye, they need to overcome the very important additional barrier represented 
by the cornea, which is the main entrance to the inner eye. The area of contact of the 
drug with the cornea is restricted to approximately 2 cm.2 This small fraction of drug 
in contact with the cornea is then confronted with the very restrictive sub-barriers 
such as the epithelium, the stroma and the endothelium. Both the first and the last 
barrier, but particularly the first, limit the absorption to water soluble substances, 
due to the existence of tight junctions between the epithelial cells.6 The stroma, with 
high water content, limits the absorption of lipophylic drugs.7 As a result of the 
above mentioned processes, typically less than 1-5% of the instilled dose reaches 
the aqueous humour.2,8 This extremely low "ocular bioavailability" often implies 
the necessity of frequent dose administration, a situation that may lead to a significant 
systemic absorption and the corresponding side effects. In some instances, the 
required posologic regimen is unviable and hence the intravitreal injection becomes 
necessary to achieve significant drug levels in the intraocular structures. 
Nanoparticular Carriers for Ocular Drug Delivery 651 
These biopharmaceutical constraints clearly evidence the necessity to conceive 
new ocular drug delivery strategies aimed at overcoming the above indicated barriers. 
Unfortunately, the requisite to preserve both the specific characteristics of 
the visual apparatus and the visual acuity, together with the inherent sensitivity of 
the eye, limit the possibilities of designing optimized ocular drug delivery systems 
substantially. 
Among the delivery strategies aimed at circumventing the above described limitations, 
the design of polymer nanoparticulate carriers offer unique features, while 
still benefit from their presentation in a liquid form. Two types of nanoparticulate 
carries have been described for ocular drug delivery: matrice-type nanoparticles, 
in which the biologically active molecule is entrapped or simply adsorbed onto 
their surface; and reservoir-type nanocapsules, which consist of a polymeric wall 
surrounding a liquid drug-containing core. Within the context of this chapter, the 
matrice-type and the reservoir-type will be termed nanoparticles and nanocapsules 
respectively. The fabrication processes of these nanostructures will not be a subject 
of description in this chapter, since they have already been reviewed.9 
Despite the above indicated limitations in the design of ocular drug delivery 
systems, the efforts oriented towards the use of nanotechnologies have been relevant 
and they have led to significant progress in the field. In this chapter, we review 
the advances made in the design of nanoparticulate carriers intended for topical 
ocular drug delivery. These nanocarriers are classified into three categories: first 
generation of basic nanoparticles and nanocapsules, second generation of nanoparticles 
and nanocapsules with a hydrophilic polymer coating, and the third generation 
of functionalized nanoparticles/nanocapsules (Fig. 1). On the other hand, being 
conscious of the fact that the progress made in this field has not yet resulted in significant 
improvements in the therapy of inner-eye diseases, the potential of nanoparticles 
as injectable ocular drug delivery vehicles will also be described in this chapter. 
Indeed, polymer nanoparticles may circumvent the problem of frequent intravitreal 
injection by providing a controlled delivery of the encapsulated drug, thus reducing 
the clinical complications associated with this modality of administration. 
2. Nanoparticulate Polymer Compositions as Topical Ocular 
Drug Delivery Systems 
As previously mentioned, the eye defense mechanisms represent the main limitation 
to the use of liquid formulations for ophthalmic therapy. Within this context, 
nanoparticles offer great possibilities of increasing the amount of drug at the anterior 
chamber of the eye, while spacing the dose administration. Table 1 and Table 2 
summarize the literature reports on the use of nanoparticulate polymer compositions 
as topical ocular drug delivery systems. 
652 Sanchez & Alonso 
First Generation 
Second Generation 
(Coating approach) 
Third Generation 
(Functionalized 
nanocarriers) 
targeting 
coating 
Fig. 1. Schematic representation of different nanosystems intended for ocular drug 
delivery: "first generation" of basic matrice-type nanoparticles (the biologically active 
molecule is entrapped or simply adsorbed onto their surface) and reservoir-type nanocapsules 
(the biologically active molecule is dissolved in a liquid core surrounded by a polymeric 
wall); "second generation" of nanoparticles and nanocapsules (the figure shows a 
nanocapsule) with a hydrophilic polymer coating; and "third generation" of surface functionalized 
nanoparticles/nanocapsules (the figure shows a nanocapsule functionalized with 
antibodies). 
2.1. First generation: Polymer nanoparticles and nanocapsules 
for topical ocular drug delivery 
Nanoparticles, primarily developed for i.v. administration, were first proposed for 
ophthalmic drug delivery in 1981. Indeed, it was Gurny and co-workers who first 
indicated the potential advantages of nanoparticles (named pseudo-latexes) over 
aqueous polymer solutions. More specifically, these authors found that pilocarpine 
adsorbed onto nanoparticles (0.3 /tm) made of cellulose acetate phthalate (CAP) 
were able to maintain a constant miosis in the rabbit for up to 10 hours, compared 
with a 4-hour response attained for pilocarpine eye drops.10 This initial report was 
followed by a number of studies aimed at evaluating the potential of different 
types of polymers including acrylic polymers, and, especially, poly (alkyl cy anoacrylates) 
(PACA), polyesters, i.e. poly-e-caprolactone, and polysaccharides, such as 
hyaluronic acid and chitosan, for ocular drug delivery. A summary of the results 
obtained with these different nanoparticulate formulations is presented in Table 1. 
Nanoparticular Carriers for Ocular Drug Delivery 653 
Table 1 Nanoparticulate compositions used in ocular drug delivery (topical 
administration). 
Polymer typea Drug 
(System type) 
In vivo resultsb (references) 
CAP 
Eudragit® 
PIPAA 
PACA 
PECL 
Pilocarpine 
(Nanoparticles) 
Ibuprofen/ 
Flurbiprofen 
(Nanoparticles) 
Cloricromene 
(Nanoparticles) 
Epinephrine 
(Nanoparticles) 
3 H-Progesterone 
(Nanoparticles) 
Pilocarpine 
(Nanoparticles) 
Betaxolol 
(Nanoparticles) 
Amikacine 
(Nanoparticles) 
3 H-Cyclosporin 
(Nanocapsules) 
Metipranolol 
(Nanocapsules) 
Betaxolol 
(Nanocapsules/ 
Nanoparticles) 
Carteolol 
(Nanocapsules/ 
Nanoparticles) 
Cyclosporin A 
(Nanocapsules) 
Indomethacin 
(Nanocapsules/ 
Nanoparticles) 
10 Prolonged miosis 
Improved "ocular bioavailability" (aqueous 
humour drug levels) and inhibition of the miosis 
induced by a surgical trauma21-22 
Improved "ocular bioavailability" (aqueous 
humour drug levels)23 
Prolonged IOP lowering effect24 
Reduced drug concentrations in cornea, 
conjunctiva and aqueous humor19 
Prolonged miosis11 
Prolonged miosis and improved reduction of IOP12 
Improved "ocular bioavailability" (aqueous 
humour drug levels), prolonged miosis and 
improved reduction of IOP for pilocarpine-loaded 
nanoparticles13 
Prolonged intraocular pressure (IOP) lowering 
effect14 
Improved "ocular bioavailability" (corneal and 
aqueous humour drug levels)16 
Prolonged therapeutic levels in cornea, sclera, uvea 
and retina as compared to those provided by an 
oily ciclosporin solution17 
Reduction of cardiovascular side effects and 
enhanced IOP lowering effect27,28 
Enhanced IOP lowering effect in a greater extent 
than PACA or PLGA. Effect more important for 
PECL nanocapsules than for nanoparticles26 
Improved IOP lowering effect, being this effect 
superior for nanocapsules than for nanoparticles. 
Reduction of cardiovascular side effects for 
nanocapsules15 
Increased and more prolonged cyclosporin corneal 
levels as compared with those corresponding to an 
oily cyclosporin solution32 
Improved bioavailability (e.g. cornea, aqueous 
humour and iris-ciliary body drug levels) as 
compared with that of indomethacin-loaded 
Microparticles (6 /xm), and that of a control 
solution30 
654 Sanchez & Alonso 
Table 1 (Continued) 
Polymer typea Drug In »i»o results" (references) 
(System type) 
Chitosan Cyclosporin A Higher and more prolonged cyclosporin levels at 
(Nanoparticles) external ocular tissues (e.g. cornea and conjunctiva) 
and negligible intraocular and systemic levels, as 
compared with those corresponding to a 
cyclosporin suspension in a chitosan 
solution47 
aCAP: Cellulose acetate phthalate; PIPAA: Poly (isopropylacrylamide); Eudragits®: Copolymers of ethylacrylate, 
methyl-methacrylate and chlorotrimethyl-ammonioethyl-methacrylate; PECL: Poly-epsiloncaprolactone; 
PLA: PolyQactic acid); PACA: Poly(alquilcyanoacrylate). 
bIOP: intraocular pressure. 
2.1.1. Acrylic polymers-based nanoparticles 
The first study reporting the potential of polyacrylic nanoparticles for ocular drug 
delivery was published by Wood et al.5 More specifically, these authors found 
that PACA nanoparticles were significantly retained in the precorneal area, and 
therefore could act as nanoreservoires for extended drug delivery. Indeed, this 
improved retention of the carrier at the ocular surface was the explanation for the 
increased drug concentration in the cornea, and for the enhanced and/or prolonged 
pharmacological effect reported for antiglaucoma drugs, such as pilocarpine,11-13 
betaxolol14 and carteolol,15 the aminoglucoside amikacine16 and the immunosuppressive 
peptide cyclosporin A.17 Moreover, some authors found that the retention 
and residence time of nanoparticles was significantly increased in inflamed ocular 
tissues.18 This was attributed to an enhanced epithelial permeability of the swollen 
conjunctival tissue. There have also been examples for which the "ocular bioavailability" 
of the drug associated to PACA nanoparticles was reduced, compared with 
that of the free drug. This negative result has been attributed to the great affinity of 
the drug (progesterone) for the polymer, and consequently, to its deficient delivery 
from the carrier.19 
Unfortunately, despite the reported efficacy of PACA nanoparticles for enhancing 
the "ocular bioavailability" of drugs, the study reported by Zimmer et al. in 
1991 evidenced that the nanoparticles entered the corneal epithelial cells, causing 
a disruption of cell membranes.20 Whether this negative result was due to specific 
experimental conditions of the reported study, or to the intrinsic nature of the PACA 
and its degradation products, remains to be clarified. Nevertheless, this could possibly 
provide the explanation for the little attention that PACA nanoparticles have 
attracted over the last decade, for this specific application. 
Nanoparticular Carriers for Ocular Drug Delivery 655 
Interestingly, the toxicity reported for PACAs has not dissuaded the research 
on the other types of acrylic polymers such as copolymers of ethylacrylate, methylmethacrylate 
and chlorotrimethyl-ammonioethyl-methacrylate (Eudragits®)21-23 
and polyacrylamide.24 In the case of Eudragit® nanoparticles, the in vivo results 
showed an enhanced bioavailability (aqueous humour drug levels) of the anti 
inflamatory drugs (e.g. ibuprofen, flurbiprofen), as well as an improved pharmacological 
response.21-23 On the other hand, using epinephrine-loaded polyacrylamide 
nanoparticles, it was possible to prolong the intraocular pressure (IOP) lowering 
effect caused by epinephrine.24 While these studies underline the positive interaction 
of acrylic nanoparticles with the ocular mucosa, further studies in terms 
of tolerance and toxicity are required to assess the practical application of these 
nanoparticles. 
2.1.2. Polyester-based nanoparticles and nanocapsules 
The results obtained for PACA nanoparticles stimulated the search for new 
nanoparticulate carriers made of different polymers. Within this context, poly-ecaprolactone 
(PECL) and the copolymers of lactic and glycolic acid (PLGA) have 
received a great deal of attention. The choice of these polymers was based on their 
broad history of safe use in humans as suture materials and implants. Moreover, at 
present, there is significant evidence of the adequate ocular tolerability of nanoparticles 
composed of these materials.25 
According to our knowledge, the first reports on the efficacy of polyester for 
topical ocular drug delivery were published in 1992.26,27 Marchal-Heussler et al.26 
compared the performance of nanoparticles made of PACA, PECL and PLGA containing 
betaxolol as a model drug, with that of the commercial eye drops. The 
results showed that under the experimental conditions assayed, all nanoparticulate 
systems yielded an improved pharmacological response (i.e. intraocular pressurelowering 
effect), compared with an aqueous eye drop control formulation of the 
drug, with the optimum responses being ascribed to PECL nanocapsules. In this 
case, the improved pharmacological effect was believed to be due to the agglomeration 
of these hydrophobic nanoparticles in the conjunctival sac, thus forming a 
depot from which the drug is slowly delivered to the precorneal area. 
Simultaneously to this work, we reported the efficacy of PECL nanocapsules 
consisting of an oily core and a PECL wall for the ocular delivery of metipranolol. 
The results of this study performed in rabbits led us to conclude that PECL 
nanocapsules were not only able to increase the pharmacological effect of this 
drug, but also able to reduce the cardiovascular side effects associated with its 
systemic absorption.27,28 This positive behavior suggested that the nanocapsules 
enhanced the drug transport across the cornea and reduced the systemic absorption 
656 Sanchez & Alonso 
through the conjunctiva. In our attempt to elucidate the mechanism of the action of 
these nanocapsules, we labeled them with rhodamine B and followed their interaction 
with the rabbit cornea and conjunctiva by the use of confocal laser scanning 
microscopy. The results of these ex vivo studies clearly evidenced that PECL 
nanocapsules were able to penetrate the corneal epithelial cells, and that they also 
exhibited a preference for the cornea vs the conjunctiva.29 Consequently, this differentiated 
interaction with both the epithelial surfaces could be taken as an adequate 
explanation for their reported ability to enhance corneal penetration and reduce systemic 
absorption. More importantly, no evidence of membrane alteration or signs 
of toxicity were detected in this study. 
With the intention of investigating the determinants of the interaction of these 
particles with the ocular mucosa, and thus of their performance as ocular drug carriers, 
we compared the efficacy of PECL nanoparticles and nanocapsules (200 nm) 
with that of PECL microparticles (6 /xm). The results clearly evidenced that the 
nanoscale size was critical with regard to the ability of the particles to enhance 
the ocular bioavailability of indomethacin30 (Fig. 2). Consequently, these results 
led us to hypothesize that nanoparticles have a greater ability than microparticles 
to interact with the corneal epithelium cells. A similar conclusion was extracted 
recently for the interaction of the polyester particles with the conjunctival cells.31 
In specific, these authors found that the in vitro uptake of nanoparticles by primary 
cultured rabbit conjunctival epithelial cells was more important than that of the 
microparticles. 
The positive results obtained for metipranolol and indomethacin encouraged 
us to test the performance of these nanosystems for the topical delivery of the 
1223 Control solution (Indocollyre) 
• Nanocapsules 
^ Emulsion 
B Nanoparticles 
I OH Microparticles 
~»rrfcn 
Fig. 2. Indomethacin concentrations attained in the aqueous humor following the topical 
application, in rabbits, of indomethacin-loaded carriers and a control drug solution (Mean 
values ± SD, n = 3,4) (*P < 0.05 compared with Indocollyre; **P < 0.05 compared with 
colloidal suspensions) (Reprinted from Ref. 30, with permission from Pharmaceutical Press.) 
Nanoparticular Carriers for Ocular Drug Delivery 657 
immunosuppressive peptide cyclosporin A. Interestingly, following topical administration 
of PECL nanocapsules containing cyclosporin A, we observed corneal 
levels of the drug which were five times higher than those provided by an oily 
solution (topical formulation of cyclosporin typically used).32 These high levels 
were not, however, translated into high drug concentrations in the aqueous humor; 
a result that was attributed to the important hydrophobicity of this peptide and its 
tendency to associate with lypophylic components. 
Therefore, at present, there is a proof-of-concept of the efficacy of polyester 
nanocapsules for enhancing the concentration of topically applied drugs in the 
corneal epithelium. Whether this enhanced concentration may or may not lead 
to a favored accumulation of the drug in the inner eye is expected to be largely 
dependent on the physicochemical characteristics of the drug. 
2.1.3. Polysaccharide-based nanoparticles 
The polyester polymers described above are hydrophobic polymers that need to 
be biodegraded into hydrophilic oligomers in order to be eliminated from the 
body. A very different class of polymers, which has only received attention in 
the last few years, is the one represented by the hydrophilic polysaccharides. 
Hyaluronic acid and chitosan are two types of polysaccharides which have opened 
new prospects in the ocular drug delivery area. The choice of hyaluronic acid has 
been justified by its bioadhesive character,33'34 but also by its well known safety 
profile. In fact, hyaluronic acid is already being used as a substitute for vitreous 
humor in intraocular surgery, since it constitutes a basic component of the vitreous 
body.35 On the other hand, chitosan is a polycationic biopolymer which exhibits 
several favorable biological properties for ocular drug administration.3 These 
properties include mucoadhesiveness,36,37 biodegradability in the rich lysozymecontaining 
mucus (i.e. ocular mucosa),38-40 and also wound healing and antimicrobial 
activity.41-43 
Despite the number of articles showing the efficacy of hyaluronic acid solutions 
for improving the retention of drugs applied topically onto the eye,33,34,44 the only 
particulate formulation that has been tested in vivo was composed of microparticles 
(1-10 ftm) rather than nanoparticles. These hyaluronate microparticles were shown 
to increase the residence time of the model drug methylprednisolone at the ocular 
surface of the rabbit eye.45 Taking into account the reported influence of the size 
on the interaction of particles with the ocular mucosa, we have recently designed 
nanoparticles consisting of hyaluronic acid and chitosan.46 At present, we know that 
these nanoparticles are stable upon incubation in simulated lachrymal fluids and 
in vivo studies are in progress in order to evaluate their mechanism of interaction 
with the ocular mucosa. 
658 Sanchez & Alonso 
Chitosan has also received significant attention in the ophthalmic field.3 One 
of the chitosan-based systems that has exhibited an interesting behavior following 
topical ocular administration, is the one consisting of chitosan nanoparticles. These 
nanoparticles have been tested on the rabbit model for their ability to enhance 
the concentration of cyclosporin A at the level of the ocular mucosa. As expected, 
the results showed that the chitosan nanoparticles were able to increase the concentrations 
of cyclosporin A in the external ocular tissues (cornea and conjunctiva) 
significantly for up to 48 hr post-instillation (Fig. 3). Despite this enhanced retention 
of the drug in the external tissues, the levels attained in the internal ocular structures 
(i.e. aqueous humor, iris and ciliary body) and in the blood were negligible. Consequently, 
these results suggested the utility of this new formulation for the treatment 
of surface eye diseases, i.e. dry eye or inflammatory diseases.47 These high drug 
concentrations restricted to the periocular tissues were later explained by a high 
corneal and conjunctival surface retention of chitosan nanoparticles. Indeed, in a 
study consisting of evaluating the concentration of fluorescent chitosan, either in the 
form of nanoparticles or as a solution, in cornea and conjunctiva, we could conclude 
that the affinity of chitosan for the ocular surface is greater when it is in a particulate 
form.48 This conclusion invites interesting prospects with regard to the potential of 
chitosan nanoparticles as drug carriers for topical ocular administration. Keeping 
this in mind, we tested the acute tolerance of chitosan nanoparticles following topical 
instillation to rabbits very recently. The results gave evidence of an excellent 
tolerance, without any sign of irritation or damage of the ocular surface structures.49 
CyA concentration in the cornea 
(ng CyA/g cornea) 
8000 
7000 
6000 
5000 
4000 
3000 
2000 
1000 
0 
2 6 24 48 
Time (h) 
Fig. 3. Cyclosporin (CyA) concentration in the cornea after topical administration in rabbits 
of CyA-loaded chitosan (CS) nanoparticles and control formulations consisting of a CyA 
suspension in a CS aqueous solution and a CyA suspension in water ^statistically significant 
differences, P < 0.05). (Reprinted from Ref. 47, with permission from Elsevier.) 
^m oyM-iuaueu uo nanupamcies 
E22 CyA suspension in a CS solution 
I I CyA suspension in water 
Nanoparticular Carriers for Ocular Drug Delivery 659 
2.2. Second nanoparticles generation: The coating approach 
The previously described nanoparticular polymer-based carriers, are shown to 
increase the intensity and contact time of drugs with the eye. Moreover, in some 
cases, this improved contact led to an enhanced intraocular penetration of drugs. 
Despite the difficulties for comparing the performance of the first-generation 
nanosystems, it is obvious that their interaction with the ocular surface is determined 
not only by the nanoscale size, but also by the surface composition of the 
nanomatrice. Taking this into account, a different approach has arisen based on 
the principle of providing to the nanocarrier, a polymer coating that favors its 
interaction with the ocular mucosa. Using this approach, it is additionally possible 
to select the adequate core composition in order to facilitate the entrapment 
and protection of the desired drug. Moreover, one can envisage the design 
of a nanocarrier with a differentiated interaction with the cornea and conjunctiva. 
An element that could be taken in consideration to achieve this aim is the 
presence of the mucus layer covering the conjuctival epithelium (i.e. where the 
goblet cells are) and its reduced amount onto the corneal surface. In this sense, 
it is important to keep in mind that the interaction with the cornea would be 
the choice for the drugs whose target is located in the inner eye. In contrast, 
the improved interaction and controlled release at the conjunctival level could 
offer a potential for the treatment of surface ocular diseases. Table 2 summarizes 
the characteristics of the different coated nanostructures developed under these 
bases. 
2.2.1. Poly a cry lie coa ting 
The first "coating approach" was intended to confer the nanosystems with a 
mucoadhesive character. Theoretically, the coating with mucoadhesive polymers 
could markedly prolong the residence time of the nanocarriers, since their clearance 
from the eye surface would be controlled by the much slower rate of mucus 
turnover than the tear turnover rate. 
The simplest approach towards this aim has been the suspension of the nanocarrier 
in an aqueous solution containing a mucoadhesive polymer. Indeed, Zimmer 
et al.50 observed that the co-administration of pilocarpine-loaded albumin nanoparticles 
with bioadhesive polymers such as poly aery lie acid (Carbopol®), hyaluronic 
acid, mucin or sodium carboximethylcellulose, led to an enhancement of the intraocular 
pressure lowering effect in rabbits. The efficacy of this approach was also tested 
for PAC A nanoparticles in an ex vivo study using bovine corneas. The results showed 
that the corneal penetration of cyclosporin A, entrapped in PACA nanoparticles, 
was improved when the nanoparticles were suspended in a polyacrylic acid gel.51 
660 Sanchez & Alonso 
Table 2 Polymer-coated nanoparticulate compositions used in ocular drug delivery (topical 
administration). 
Polymer coating3 Coreb 
composition 
Associated In vivo results (references) 
drug/marker 
Polyacrylic acid Albumin Pilocarpine Enhanced intraocular pressure 
Nanoparticles lowering effect and duration of 
PECL/oil Indomethacin Improved drug "ocular 
Nanocapsules bioavailability" (corneal and 
aqueous humor drug levels)20 
PECL/oil Rhodamine Enhanced retention of the 
Nanocapsules nanocapsules on the ocular 
surface54 
PECL — Not reported52 
nanoparticles 
PACA Acyclovir Improved drug "ocular 
Nanoparticles bioavailability" (aqueous humour 
drug levels)59 
PLA Acyclovir Improved drug "ocular 
Nanoparticles bioavailability" (aqueous humour 
drug levels)25 
PECL Rhodamine Evidence of the ability of 
nanocapsules PEG-coated nanocapsules to cross 
the corneal epithelium layers54 
Chitosan 
Chitosan 
Hyaluronic acid 
PEG 
PEG 
PEG 
aPEG: Poly(ethyleneglycol). 
bPECL: Poly-epsilon-caprolactone; PLA: Polyflactic acid); PACA: Poly(alquilcyanoacrylate). 
2.2.2. Polysaccharide coating 
As indicated in the previous section covering the nanocarriers of first generation, 
two polysaccharides have attracted special attention as mucoadhesive materials 
for ocular application: hyaluronic acid and chitosan. Apart from the simple dispersion 
of the core material into an aqueous polymer solution described above, the 
first attempt to efficiently coat nanoparticles with hyaluronic acid was described 
by Barbault-Foucher et al.52 These authors described different strategies for the formation 
of hyaluronate-coated poly-e-caprolactone (PECL) nanoparticles intended 
for ocular drug delivery. These strategies were simple adsorption, ionic promoted 
interaction and chain entanglement during the nanoparticles fabrication process. 
While the in vivo efficacy of these strategies remains to be investigated, this publication 
shows the versatility of the coating approach procedure. 
The mucoadhesive polysaccharide chitosan has also been identified as a successful 
candidate for the "coating approach". The mucoadhesive properties of 
Nanoparticular Carriers for Ocular Drug Delivery 661 
chitosan have generally been ascribed to its polycationic nature, which promotes 
the interaction with the negatively charged ocular mucosa. However, the cationic 
nature should not be taken as the only factor determinant of the mucoadhesive 
properties of polymer-coated nanocarriers. In fact, in a previous study, we have 
shown that the performance of PECL nanoparticles coated with two different 
polycationic polymers (poly-L-lysine and chitosan) was drastically different. Concretely, 
we observed that PECL nanoparticles coated with chitosan were significantly 
more efficient at increasing the corneal uptake of the encapsulated molecule 
(14C-indomethacin) in rabbits, than those coated with poly-L-lysine.53 Therefore, 
these results led us to conclude that it was the intrinsic mucoadhesive character 
of chitosan, not exclusively ascribed to its positive charge, that is the reason for its 
successful behavior. 
More recently, we attempted to investigate ex vivo (isolated rabbit cornea) and 
in vivo the mechanism of interaction of chitosan-coated PECL nanocapsules with 
the cornea.54 The results of this study showed that rhodamine encapsulated in 
these systems had an improved transport across the cornea, compared with that of 
the marker alone, or in combination with blank nanocapsules (Fig. 4). Moreover, 
the examination of the corneas treated with fluorescent nanocapsules by confocal 
microscopy suggested that CS-coated nanocapsules have a lower penetration 
oi 300 
CD 
c 
o o 
0) 
x: 
 
w 
2 
o 
(0 
•o 
d) 
•e o 
Q.  
c 
2 
+ J 
200 
100 
• Free Rd 
E3 Free Rd + blank CS-coated nanocapsules 
B Rd-loaded CS-coated nanocapsules 
Time (h) 
Fig. 4. Rhodamine (Rd) amount transported across the rabbit cornea in an ex vivo study 
for Rd-loaded chitosan (CS) coated nanocapsules and control formulations consisting of a 
Rd solution and a physical mixture of free Rd and blank CS-coated nanocapsules ^Statistically 
significant different from free Rd). ("^Statistically significant different from free Rd and 
free Rd plus blank CS-coated nanocapsules) (Reprinted from Ref. 54, with permission from 
Elsevier). 
662 Sanchez & Alonso 
extent than the non-coated PECL nanocapsules, a result that could be attributed 
to the increased surface retention of the nanocapsules in the mucus layer. Therefore, 
overall, the results obtained until now with nanoparticulate carriers coated 
with mucoadhesive polymers permit us to conclude the efficacy of this approach, 
in terms of increasing the retention of the nanoparticles at the eye surface. This 
improved retention could be translated depending on the solubility properties of 
the drug encapsulated to a more important corneal penetration, or in a greater 
retention on the ocular surface. 
2.2.3. Polyethyleneglycol (PEC) coating 
A very different alternative in the "coating approach" has been the one intended to 
provide the nanoparticulate carrier with an improved stability upon contact with 
the mucosal fluids. In fact, both the mucus layer and the lachrymal fluids are very 
rich in enzymes and proteins, which may be attached to the nanoparticles and 
promote their aggregation.55 Poly(ethylene glycol) (PEG) appears to be an ideal 
candidate for such purpose, since it has been widely used to prevent the interaction 
of colloidal carriers with proteins. For example, in a study performed by 
us,56 we observed that a PEG coating around PLA nanoparticles prevented their 
aggregation in the presence of lysozyme (highly concentrated in the mucus layer). 
On the other hand, this stabilizing effect has been the main explanation for the 
successful behavior of PEG-coated PLA nanoparticles as transmucosal drug carriers 
(i.e. after nasal and oral administration).57,58 Therefore, from these studies we 
suggested that the presence of a hydrophilic PEG layer onto the surface of polyester 
nanoparticles could result in an enhanced stability and, hence, to an improved interaction 
of these nanosystems with the ocular mucosa. 
The first report on the positive effect of the PEG coating approach on ocular drug 
administration was published by Fresta et al.59 These authors evaluated the ocular 
bioavailability of acyclovir-loaded PEG-coated PACA nanoparticles and observed 
a significant increase of the drug levels in the aqueous humor, when comparing 
these systems with an aqueous acyclovir suspension and with a physical mixture 
of the free drug and the blank PEG-coated PACA nanoparticles. Interestingly, in this 
work, the authors did not consider the possibility of an improved stability. Rather, 
they suggested that the PEG-coated particles might have an improved mucoadhesion. 
However, no studies were reported to verify this mechanistic hypothesis. In a 
more recent work, the same group reported the efficacy of the PEG coating but with 
a different core (PLA nanoparticles), in terms of increasing the "ocular bioavailability" 
(aqueous humor drug levels) of acyclovir.25 Moreover, these authors observed 
that the positive effect of the PEG coating disappeared, when the mucus layer was 
removed using N-acetylcysteine, prior to the administration of these systems to 
Nanoparticular Carriers for Ocular Drug Delivery 663 
rabbits. This observation led the authors to suggest that the mucoadhesion of the 
PEG-coated nanocarriers may play a role in its mechanism of action. However, 
this result could also be understood as a consequence of the stabilizing effect provided 
by PEG in the mucosal surfaces; this effect being not visible in the absence 
of mucus. 
In order to obtain further insights into the interaction of PEG-coated nanoparticles 
with the cornea, we have recently compared their behavior with that of the 
uncoated PECL nanocapsules and the chitosan-coated nanocapsules. The confocal 
images showed that the three types of nanocapsules were able to enter the 
corneal epithelium. However, their penetration depth followed the ranking of PEGcoated 
nanocapsules > uncoated nanocapsules > chitosan-coated nanocapsules.54 
The more important corneal penetration of PEG-coated nanocapsules, as compared 
with that of the non-coated ones, was suggested to be a consequence of their 
improved stability in the mucosal fluids. The chitosan coating is also known to 
affect the stability of colloidal particles positively in the presence of the proteins 
such as lysozyme.56 However, in this case, the superficial retention of chitosancoated 
nanocapsules (described above) could also be understood as a result of their 
mucoadhesive character. 
Overall, the results obtained so far with nanoparticulate drug carriers, coated 
with hydrophilic polymers, indicate that depending on their nature, these polymers 
are able to increase the stability and/or the mucoadhesion of the nanocarrier. It 
could also be presumed that an increase in the mucoadhesion should lead to an 
accumulation of the drug carrier at the ocular surface. Also, in the case of drugs with 
adequate permeability properties, it should lead to an increase and prolongation in 
the corneal penetration. Similarly, an increase in the stability is expected to lead to 
a more important interaction and transport of the nanoparticles across the corneal 
epithelium. Moreover, these results suggest that both the extent of the interaction 
and the penetration depth of the nanocarriers with the cornea, can be modulated 
by providing them with an appropriate coating. Despite the need of additional 
mechanistic studies as evidence, the results reported so far provided some basis 
for the development of strategies intended as an efficient drug targeting to specific 
ocular structures, as discussed below. 
2.3. Third nanoparticles generation: Towards 
functionalized nanocarriers 
The new tendency in the design of the new drug delivery systems is directed 
towards integrating several drug delivery technologies, in order to provide the 
system with unique properties. In the particular case of the ocular drug delivery, 
the design of highly sophisticated drug delivery nanosystems could benefit from 
664 Sanchez & Alonso 
the knowledge gained from the application of such systems to other trasmucosal 
routes of administration. As described in the previous chapters of this book, some of 
the present efforts on the design of more specialized nanocarriers go through functionalizing 
their surface. This means that the design of nanoparticles with surface 
characteristics allowing their functionalization with specific targeting moieties, is 
able to selectively direct the nanocarrier to the predetermined ocular structures. 
Among the targeting moieties described till now, lectins may represent an interesting 
option for targeting the ocular mucosa. Lectins are glycoproteins capable of 
recognizing and binding reversibly to specific carbohydrate moieties which are 
present on cell surfaces and mucin. In fact, lectin-like molecules are known to be 
important in the adhesion of micro-organisms to mucosal surfaces.60 Therefore, 
lectins clearly differ from conventional mucoadhesive materials which interact nonspecifically 
with the mucus or simply adhere to biological surfaces.61 Some examples 
of lectins are wheat germ agglutinin and phasoleus vulgaris agglutinin, which 
bind specifically to N-acetylgalactosamine and mannose receptors, respectively. 
With regard to the potential of lectins as targeting moieties for ocular drug 
delivery, it is a known fact that lectins can bind to the corneal and conjunctival 
surfaces and also to some constituents of the tear film.62'63 From our knowledge, 
despite this information, there is no evidence of the potential of the targeted systems 
in the ocular drug delivery field. 
A different category of targeting ligands is represented by the monoclonal 
antibodies. The initial attempts towards the monoclonal antibody-based targeting 
approach have been directed to the treatment of ocular herpes simplex virus 
(HSV) infection. More specifically, Norley et al.M proposed the attachment of monoclonal 
antibodies (anti-glycoprotein D of HSV) to liposomes, in order to achieve 
the targeted delivery of antiviral drugs to the infected corneal cells. The results of 
this work showed the ability of these "immunoliposomes" to preferentially bind to 
virus-infected corneal cells in vitro. However, there are no in vivo data available to 
support the efficacy of this targeting strategy, thus far. 
The lack of reported success of the monoclonal antibody-based targeting 
approach could be related to the late clinical development of antibodies. Nevertheless, 
the enormous effort devoted to the development of antibodies for therapeutic 
or diagnosis purposes in the last few years65 opens optimistic prospects with regard 
to their use as targeting moieties for nanoparticulate carriers. 
Within this context, PEG-coated nanoparticles offer interesting opportunities 
for the functionalization with ligands such as lectins66 and monoclonal antibodies.67 
Additionally, as polysaccharides present many available reactive groups, active 
targeting could also be attained by grafting ligands onto the polysaccharide-coated 
nanoparticles. 
Nanoparticular Carriers for Ocular Drug Delivery 665 
3. Nanoparticulate Polymer Compositions as Subconjuctival 
Drug Delivery Systems 
The subconjunctival route has been proposed as an alternative to the topical drug 
delivery route, in order to force the retention of a significant amount of drug in 
the eye. The drug molecules locate underneath the conjunctival epithelium are 
supposed to diffuse through the sclera and reach the inner eye. There is no doubt 
that the most important limitation of this modality of administration is the poor 
acceptability by the patients. Therefore, the use of controlled release micro and 
nanoparticles was thought to be a good approach in order to reduce the number of 
injections. Despite the logic of this approach, no improved pharmacological and/or 
therapeutical effects have been reported so far for either micro or nanoparticles. For 
example, the subconjunctival injection of cyclosporin-loaded PLGA microparticles 
failed to improve the response of this drug.68 On the other hand, to the best of 
our knowledge, there has been no report on the efficacy of nanoparticles for the 
delivery of drugs at the subconjunctival level. The study reported by Amrite et al.69 
showed that the model fluorescent nanoparticles (20 nm) and microparticles (2 ^m) 
administered subconjunctival^ were not able to cross the sclera, remaining at the 
injection site. 
4. Nanoparticulate Polymer Compositions as Intravitreal 
Drug Delivery Systems 
Most diseases affecting the posterior segment of the eye are chronic in nature and 
require prolonged drug administration. These diseases are one of the major causes 
of blindness in the developed world. Unfortunately, the described difficulty of 
reaching effective drug levels at intraocular structures represents a major limitation 
associated to these therapies (i.e. treatment of proliferative vitreoretinopathy, 
endophthalmitis, recurrent uveitis, acute retinal necrosis, choroidal neovascularization 
and cytomegalovirus retinitis). In these severe situations, the intravitreal 
injection becomes the route of choice for drug delivery. However, in clinical practice, 
this modality of administration has important draw-backs: (i) poor patient 
acceptability, which may lead to failure of therapy; (ii) rapid drug elimination from 
vitreous humor (i.e. removal to the systemic circulation along with the aqueous 
humor drainage, active secretion from the retina); (iii) possible retinal toxicity of 
certain potent drugs; (iv) potential hazards associated with repeated intravitreal 
injection, such as the clouding of the vitreous humor, retinal detachment, lens damage 
and endophthalmitis. 
666 Sanchez & Alonso 
The above indicated problems illustrate the need for the design of adequate 
controlled release systems which could minimize the frequency of injection. 
Vitrasert® is an example of a commercially available sustained release intraocular 
device for ganciclovir, which has been approved for use in patients suffering from 
cytomegalovirus (CMV). This implant is a reservoir system consisting of a magnesium 
stearate core containing the drug, and a coating of ethylenevinyl acetate polymers. 
Apart from the necessity of surgical removal,70 additional problems observed 
for this device include endophthalmitis, retinal detachment, dislocation of implant 
and poor intravitreal drug levels due to its placement in the suprachoroidal space.71 
Within this context, biodegradable micro and nanoparticles appear to offer 
advantages when compared with large devices, since they can be injected through 
a needle, thus avoiding the necessity of a surgical procedure. Among the particulate 
carriers investigated to date for intraocular drug delivery, those made of 
biodegradable polyesters such as poly (lactic/gly colic acid) (PLGA) are expected 
to offer a significant potential. In fact, there are already a number of reports on the 
biodegradability, tolerability and efficacy of PLGA intraocular implants72,73 and 
microparticles (for a review, see Ref. 74). 
With regard to the specific potential of nanoparticles, in two studies published 
in 199475'76 aimed at evaluating the interaction of PLA microparticles (0.2-1 /zm) 
with retinal pigment epithelium cells, it was found that these particles were internalized 
by the above mentioned cells. This finding was justified by the known 
phagocytic activity of these cells, which, on the other hand, are essential for the 
maintenance of retinal metabolism and visual acuity. This initial observation was 
corroborated in a more recent study which evidenced that PLGA nanoparticles 
were localized within these cells even at 4 months post-administration.77 Moreover, 
these authors observed that PLGA nanoparticles were well tolerated following 
intravitreal injection to rats. Therefore, these publications showed the potential of 
nanoparticles no only as simple depot controlled release systems but as intracellular 
controlled delivery systems for bioactive molecules. 
Surprisingly, despite the attractive features of PLGA nanoparticles as intraocular 
delivery systems, the information reporting the success of this approach is 
scarce. For example, in a very recent publication it was shown that PLGA nanoparticles 
could work as gene delivery systems to the posterior segment of the eye.78 
Concretely, the plasmid encoding the red nuclear fluorescent protein (RNFP) was 
associated to PLGA nanoparticles and then injected into the vitreous cavity of rabbits. 
The results showed an important level of RNFP expression within the retinal 
pigment epithelial cells, thus indicating the adequate internalization and delivery 
of the plasmid into the cells. On the basis of these findings, these authors suggested 
the potential of nanoparticles for designing future gene-based ocular therapy 
strategies. 
Nanoparticular Carriers for Ocular Drug Delivery 667 
Another type of nanoparticles that has been investigated for intravitreal drug 
delivery is the one consisting of PACA. More specifically, these nanoparticles were 
tested for their ability to deliver 3H-acyclovir and ganciclovir for extended periods 
of time following intravitreal injection to rabbits. The drug concentrations attained 
in the vitreous and retina were high and steady for up to 10 days.79 Unfortunately, 
these positive results were counteracted by the negative reaction observed in the 
lens (opacification) and in the aqueous humour (turbidity). 
With respect to the alteration of the normal physiological conditions of the eye, 
one of the problems that could be expected from the use of micro and nanospheres 
is their instability in the vitreous humor. Indeed, although no stability study has 
been reported, it could be accepted that, as in the case of other biological fluids, 
i.e. lachrymal fluid, nanoparticles may suffer an aggregation process mediated 
by their interaction with proteins. One of the alternatives to resolve this problem 
could be the PEG coating approach described in the previous sections. Interestingly, 
while this approach has not been applied to PLGA nanoparticles yet, some evidence 
of its efficacy has been reported for PEG coated-poly(hexadecyl cyanoacrylate) 
nanoparticles.80 These PEG-coated nanoparticles, containing tamoxifen, have 
shown promising results for the treatment of experimental autoimmune uveitis, 
although no comparison was made between PEG coated and non-coated nanoparticles. 
Therefore, and in spite of these promising results, the potential application of 
these nanoparticles will greatly depend on their tolerability and biodegradability 
in the ocular environment. 
Finally, non-polymeric nanoparticles have also been reported for intraocular 
drug delivery. Concretely, Merodio et al.sl evaluated the ocular toxicity induced 
by the prolonged presence of the ganciclovir-loaded albumin nanoparticles after 
their intravitreal injection to rats. These authors detected the presence of these 
systems in the vitreous cavity for up to two weeks after their intraocular injection. 
In addition, according to the authors, the histological evaluation of these adjacent 
tissues revealed a good tolerance. 
In summary, the reports of the potential of nanoparticles as intraocular drug 
delivery systems indicate that while their characteristics appear to be appealing 
for such application, further studies are necessary to assess important issues that 
include their stability and biodistribution in the intraocular cavity, as well as their 
biocompatibility and absence of toxic reactions or alterations of the normal function 
of the eye. 
5. Conclusions and Outlook 
Despite extensive research in the field, the major problem in ocular drug delivery 
is the attainment of an optimal drug concentration at the intended site of action for 
668 Sanchez & Alonso 
a sufficient period of time. The site of action maybe located on the eye surface or 
in the inner ocular structures. The important barriers that need to be overcome in 
order to reach the target site limits not only the number of medications available for 
the treatment of ocular diseases, but also the extent to which those available can be 
used without incurring undesirable systemic side effects. From the results described 
in this chapter, it is possible to conclude that nanoparticles offer great chances of 
solving these limitations, while still benefiting from their topical administration as 
eye drops. Indeed, nanoparticles, depending on their composition, are significantly 
retained on the ocular mucosa, and from this location, they deliver the associated 
drugs for extended periods of time. This situation normally results in an enhanced 
and prolonged therapeutic response, and also in a decrease in the side effects. The 
results reported so far have also evidenced that both the extent of interaction and 
the penetration depth of the colloidal systems with the cornea, can be modulated 
by the selection of an appropriate coating. In addition to these beneficial effects 
associated with the topical ocular administration, nanoparticles offer an interesting 
potential in terms of improving intraocular drug administration. This potential 
includes not only the prolongation of the residence time of drugs in the eye, but 
also their targeting to the retinal cells. 
Finally, significant efforts are currently underway to develop highly sophisticated 
nanoparticles functionalized with specific targeting ligands (i.e. lectins and 
antibodies). Advances in this area are expected to open new avenues for the diagnostic 
and therapy (including gene therapy) of ocular disorders. 
Acknowledgments 
The authors would like to thank the Spanish Ministery of Education and Science 
for the financial support of some of the studies described in this chapter 
(Refs. MAT2004-04792-C02-02 and SAF2004-08319-C02-01). 
References 
1. Mitra AK (1993) Mucoadhesive polymers in ophthalmic drug delivery systems, in Mitra 
AK (ed.), Ophthalmic Drug Delivery Systems. Marcel Dekker: New York. 1-3. 
2. Lang JC (1995) Ocular drug delivery conventional ocular formulations. Adv Drug Del 
Rev 16:3943. 
3. Alonso MJ and Sanchez A (2003) The potential of chitosan in ocular drug delivery. 
/ Pharm Pharmacol 55:1451-1463. 
4. Chrai SS and Robinson JR (1973) Ocular evaluation of methylcellulose vehicle in rabbits. 
J Pharm Sci 62:1112-1121. 
5. Wood RW, Lee VHK, Kreuter J and Robinson JR (1985) Ocular disposition of poly-hexyl- 
2-cyanoacrylate nanoparticles in the albino rabbit. Int J Pharm 23:175-183. 
Nanoparticular Carriers for Ocular Drug Delivery 669 
6. Maren TH and Jankowska L (1985) Ocular pharmacology of sulfonamides: The cornea 
as barrier and depot. Curr Eye Res 4:399-408. 
7. Huang HS, Schoenwald RD and Lach JL (1983) Corneal penetration behavior of betablocking 
agents I: Physiochemical factors. / Pharm Sci 72:1272-1279. 
8. Felt O, Carrel A, Baehni P, Buri P and Gurny RJ (2000) Chitosan as tear substitute. Ocular 
Pharmacol 16:261-270. 
9. Alonso MJ (1996) Nanoparticulate drug Carrier Technology, in Cohen S and Bernstein H 
(eds.), Microparticulate Systems for the Delivery of Proteins and Vaccines. Marcel Dekker Inc., 
NewYork, pp. 203-242. 
10. Gurny R (1981) Preliminary study of prolonged acting drug delivery systems for the 
treatment of glaucoma. Pharma Acta Helv 56:130. 
11. Harmia T, Kreuter J, Speiser S, Boye T, Gurny R and Kubis A (1986) Enhancement of the 
myotic response of rabbits with pilocarpine-loaded polybutylcyanoacrylate nanoparticles. 
Int J Pharm 33:187-193. 
12. Diepold R, Kreuter J, Himmer J, Gurny R, Lee VHL, Robinson JR and Saettone MF (1989) 
Comparison of different models for the testing of pilocarpine eyedrops. Graefe's Arch Clin 
Exp Ophthalmol 227:188. 
13. Zimmer A, Mutschler E, Lambrecht G, Mayer D and Kreuter J (1994) Pharmacokinetic 
and pharmacodynamic aspects of an ophthalmic pilocarpine nanoparticle delivery 
system. Pharm Res 11:1435. 
14. Marchal-Haussler L, Maincent P, Hoffman M, Spittler J and Couvreur P (1990) Antiglaucomatous 
actyvity of betaxolol chloridrate sorbed onto different nanoparticle preparations. 
Int} Pharm 58:115-122. 
15. Marchal-Haussler L, Sirbart D, Hoffman M and Maincent P (1993) Polyepsoncaprolactone 
nanocapsules in carteolol ophthalmic delivery. Pharm Res 10:386-390. 
16. Losa C, Calvo P, Castro E, Vila-Jato JL and Alonso MJ (1991) Improvement of ocular 
penetration of amikacin sulfate by association to nanoparticles. / Pharm Pharmacol 
43:548-552. 
17. Bonduelle S, Carrier M, Pimienta C, Benoit JP and Lenaerts V (1996) Tissue concentration 
of nanoencapsulated radiolabelled cyclosporin. Eur]Pharm Biopharm 42:313-319. 
18. Diepold R, Kreuter J, Guggenbuhl P and Robinson JR (1989) Distribution of polyhexylcyanoacrylate 
nanoparticles in healthy and cronically inflammed rabbit eyes Int J Pharm 
54:149. 
19. Li VH, Wood RW, Kreuter J, Harmia T and Robinson JR (1986) Ocular drug delivery of 
progesterone using nanoparticles. / Microencapsulation 3:213-218. 
20. Zimmer A, Kreuter J and Robinson JKJ (1991) Studies on the transport pathway of PBCA 
nanoparticles in ocular tissues. Microencapsulation 8:497-504. 
21. Pignatello R, Bucolo C, Ferrara P, Maltese A, Puleo A and Puglisi G (2002) Eudragit RSI 00 
nanosuspensions for the ophthalmic controlled delivery of ibuprofen. Eur J Pharm Sci 
16:53-61. 
22. Pignatello R, Bucolo C, Spedalieri G, Maltese A and Puglisi G (2002) Flurbiprofenloaded 
acrylate polymer nanosuspensions for ophthalmic application. Biomaterials 
23:3247-3255. 
670 Sanchez & Alonso 
23. Bucolo C, Maltese A, Maugeri F, Busa B, Puglisi G and Pignatello R (2004) Eudragit RL100 
nanoparticle system for the ophthalmic delivery of cloricromene. / Pharm Pharmacol 
56:841-846. 
24. Hsiue GH, Hsu SH, Yang CC, Lee SH and Yang IK (2002) Preparation of controlled 
release ophthalmic drops. Biomaterials 23:457-462. 
25. Giannavola C, Bucolo C, Maltese A, Paolino D, Vandelli MA, Puglisi G, Lee VHL and 
Fresta M (2003) Influence of preparation conditions on acyclovir-loaded nanospheres. 
Pharm Res 20:584. 
26. Marchal-Haussler L, Fessi H, Devissaget JP, Hoffman M and Maincent P (1992) Colloidal 
drug carrier systems for the eye. STP J Pharm Sci 2:98. 
27. Losa C, Alonso MJ, Vila JL, Orallo F, Martinez J, Saavedra J A and Pastor JC (1992) Reduction 
of cardiovascular side effects associated with ocular administration of metipranolol. 
/ Ocul Pharmacol 8:191. 
28. Losa C, Marchal-Heussler L, Orallo F, Vila-Jato JL and Alonso M (1993) Design of new 
formulations for topical ocular administration. / Pharm Res 10:80-87. 
29. Calvo P, Thomas C, Alonso MJ, Vila Jato JL and Robinson J (1994) Study of the interaction 
of nanocapsules with the cornea. Int J Pharm 103:283-291. 
30. Calvo P, Alonso MJ, Vila-Jato JL and Robinson JR (1996) Improved ocular bioavailability 
of indomethacin. / Pharm Pharmacol 48:1147-1152. 
31. Qaddoumi MG, Ueda H, Yang J, Davda J, Labhasetwar V and Lee VHL (2004) The 
characteristics and mechanisms of uptake of PLGA nanoparticles in rabbit conjunctival 
epithelial cell layers. Pharm Res 21:641-648. 
32. Calvo P, Sanchez A, Martinez J, Lomez MI, Calonge M, Pastor JC and Alonso MJ (1996) 
Polyester nanocapsules as new topical ocular delivery systems. Pharm Res 13:311-315. 
33. Saettone MF, Chetoni P, Torracca MT, Burgalassi S and Giannaccini B (1989) Evaluation of 
mucoadhesive properties and in vivo active of ophthalmic vehicles based on hyaluronic 
acid. Int} Pharm 51:203-212. 
34. Saettone MF, Giannaccini B, Chetoni P, Torracca MT and Monti D (1991) Evaluation of 
sodium hialuronate as adjuvants for topical ophthalmic vehicles, hit] Pharm 72:131-139. 
35. Lutjen-Drecoll E, Schenholm M, Tamm E and Tengblad A (1990) Visualization of 
hyaluronic acid in the anterior segment of rabbit and monkey eyes. Exp Eye Res 
51:55-63. 
36. Lehr CM, Bowstra JA, Schacht EH and Junginger HE (1992) In vitro evaluation of 
mucoadhesive properties of chitosan. Int ] Pharm 78:43^18. 
37. Borchard G, Lueben HL, De Boer GA, Coos Verhoef J, Lehr CM and Junginger HE 
(1996) The potential of mucoadhesive polymers in enhancing intestinal peptide drug 
absorption. Ill: Effects of chitosan-glutamate and carbomer on epithelial tight junctions 
in vitro. J Control Rel 39:131-138. 
38. Muzzarelli RAA (1993) Chitins and chitosan. Carbohydrate Polym 20:7-16. 
39. Muzzarelli RAA (1997) Human enzymatic activities related to the therapeutic administration 
of chitin derivatives. Cell Mol Life Sci 53:131-140. 
40. Nordveit RJ, Varum KM and Smidsrod O (1994) Degradation of partially N-acetylated 
chitosans with hen egg white and human lysozyme. Carbohydrate Polym 24:253-260. 
Nanoparticular Carriers for Ocular Drug Delivery 671 
41. Balassa LL and Prudden JF (1978) The discovery of a potent pure chemical woundhealing 
accelerator. Proceedings of the First International Conference on Chitin/Chitosan 
296-305. 
42. Allan CR and Hadwiger LA (1979) The fungicidal effect of chitosan on fungi of varying 
cell wall composition. Exp Micol 3:285-287. 
43. Muzzarelli RAA (1983) Chitin derivatives. Carbohydrate Polymers 3:53-75. 
44. Camber O and Edman P (1989) Sodium hyaluronate as an ophthalmic vehicle. Curr Eye 
Res 8:563-567. 
45. Kyyronen K, Hume L, Urtti A, Topp E and Stella V (1992) Methylprednisolone esters of 
hyaluronic acid in ophthatlic drug delivery. Int J Pharm 80:161. 
46. De la Fuente M, Seijo B and Alonso MJ (2004) Nanoparticles for ophthalmic drug administration. 
PSW2004,2nd Pharmaceutical Sciences World Congress, Kyoto, Japan, p. 205. 
47. De Campos A, Sanchez A and Alonso MJ (2001) Chitosan nanoparticles: A new ophthalmic 
vehicle. Int} Pharm 224:159-168. 
48. De Campos A, Diebold Y, Carvalho ELS, Sanchez A and Alonso MJ (2002) Chitosan 
nanoparticles as new ocular drug delivery system. Pharm Res 21:803-810. 
49. Diebold Y, Salamanca AE, Jarrin M, Calonge M, Vila A, Carvalho ELS, Fuente M, 
Seijo B and Alonso MJ (2005) Nanotechnologies for ocular surface disorders. Ocular 
Surface 3:S56. 
50. Zimmer AK, Chetoni P, Saettone MF, Zerbe H and Kreuter J (1995) Evaluation of 
pilocarpine-loaded albumin particles. / Control Rel 33:31-46. 
51. Le Bourlais C, Acar L, Zia H, Sado PA, Needham T and Leverge R (1998) Ophthalmic 
drug delivery systems. Prog Retin Eye Res 17:33-58. 
52. Barbault-Foucher S, Gref R, Russo P, Guechot J and Bochot A (2002) Design of poly-Ecaprolactone 
nanospheres. / Control Rel 83:365-375. 
53. Calvo P, Vila-Jato JL and Alonso MJ (1997) Comparative in vitro evaluation of several 
colloidal systems. Int J Pharm 153:41-50. 
54. De Campos AM, Sanchez A, Gref R, Calvo P and Alonso MJ (2003) The effect of a PEG 
versus a chitosan coating on the interaction of drug colloidal carriers with the ocular 
mucosa. Eur } Pharm Sci 20:73-81. 
55. Rohen JW and Lutjen-Drecoll E (1992) Functional morphology of the conjunctiva, in 
Lemp MA, Marquard R (eds.), The dry eye: A comprehensive guide. Springer-Verlag, Berlin, 
pp. 35-63. 
56. Vila A, Sanchez A, Tobio M, Calvo P and Alonso MJ (2002) Design of biodegradable 
particles. / Control Rel 78:15-24. 
57. Tobio M, Gref R, Sanchez A, Langer R and Alonso MJ (1998) Stealth PLA-PEG nanoparticles. 
Pharm Res 15:270-275. 
58. Tobio M, Sanchez A, Vila A, Soriano I, Evora C, Vila-Jato JL and Alonso MJ (2000) The 
role of PEG on the stability in digestive fluids and in vivo fate of PEG-PLA nanoparticles 
following oral administration. Coll SurfB: Biointerfaces 18:315-323. 
59. Fresta M, Fontana G, Bucolo C, Cavallaro G, Giammona G and Puglisi G (2001) Ocular 
tolerability and in vivo bioavailability of PEG coated nanospheres. / Pharm Sci 
90:288-297. 
672 Sanchez & Alonso 
60. Calderone R and Wadsworth E (1993) Adherence molecules of Candida albicans: Analysis 
of host-pathogen interactions. Implications for pathogenesis. / Microbiol Methods 
18:197-211. 
61. Kompella UB and Lee VHL (1992) Means to enhance penetration. 4. Delivery systems 
for the enhancement of peptide and protein drugs: Design considerations. Adv Drug Del 
Rev 8:115-162. 
62. Bishop PN, Bonshek RE, Jones CPJ, Ridway AEA and Stoddart RW (1991) Lectin binding 
sites in normal, scarred, and lattice dystrophy corneas. Br / Ophthalmol 75:22-27. 
63. Nicholls TJ, Green KL, Rogers DJ, Cook JD, Wolowacz S and Smart JD (1996) Lectins in 
ocular drug delivery. Int} Pharm 138:175-183. 
64. Norley SG, Huang L and Rouse BT (1986) Targeting of drug loaded immunoliposomes. 
} Immunol 136:681-685. 
65. Walsh G (2000) Biopharmaceutical benchmarks. Nat Biotechnol 18:831-833. 
66. Gref R, Minamitake Y, Peracchia MT, Trubetskoy V, Torchilin V and Langer R (1994) 
Biodegradable long-circulating polymeric nanospheres. Science 263:1600-1603. 
67. Olivier JC, Huertas R, Lee HJ, Calon F and Pardridge WM (2002) Synthesis of Pegylated 
Immunonanoparticles. Pharm Res 19:1137-1143. 
68. Silva MVR, Rodriguez-Ares MT, Sanchez-Salorio M, Diaz MJL, Alvarez JC, Jato JLV 
and Tome CC (1999) Efficacy of subconjunctival cyclosporin-containing microspheres 
on keratoplasty rejection in the rabbit. Graefe's Arch Clin Exper Ophthalm 237:10. 
69. Amrite AC, Ayalasomayajula SP and Kompella UB (2003) Ocular distribution of intact 
nano and microparticles. Drug Del Tech 3:62-67. 
70. MacCumber MW, Sadeghi S, Cohen JA and Deutsh TA (1999) Suture loop to aid in 
ganciclovir implant removal. Arch Ophthalmol 117:1250-1254. 
71. Washington N, Washington C and Wilson C (2001) Ocular drug delivery, in 
Washington N, Washington C and Wilson C (eds.), Physiological Pharmaceutics: Barriers 
to Drug Absorption, 2nd ed. Taylor and Francis, London, 249-270. 
72. Kimura H, Hashizoe M, Nishiwaki H, Honda Y and Ikada Y (1994) Advances in ocular 
therapeutics. Invest Ophthalmol Visual Sci 35:2815-2819. 
73. Kunou NOY, Hashizoe M, Honda Y, Hyon SH and Ikada Y (1995) Controlled intraocular 
delivery of ganciclovir with use of bio-degradable scleral implant in rabbits. / Control 
Rel 37:143-150. 
74. Herrero-Vanrell R and Refojo MF (2001) Biodegradable microspheres for vitreoretinal 
drug delivery. Adv Drug Del Rev 52:5-16. 
75. Moritera T, Ogura Y, Yoshimura S, Kuriyama S, Honda T, Tabata T and Yoshito I 
(1994) Feasibility of drug targeting to the retinal pigment epithelium. Curr Eye Res 
13:171. 
76. Kimura H, Ogura Y, Moritera T, Honda, Y, Tabata Y and Ikada Y (1994) In vitro phagocytosis 
of polylactide microspheres by retinal pigment epithelial cells. Curr Eye Res 13: 
353-360. 
77. Bourges JL, Gautier SE, Delie F, Bejjani RA, Jeanny JC, Gurny R, Benezra D and Behar 
Cohen FF (2003) Ocular drug delivery targeting the retina. Invest Ophthalmol Vis Sci 
44:3562-3569. 
Nanoparticular Carriers for Ocular Drug Delivery 673 
78. Bejjani RA, Benezra D, Cohen H, Rieger J, Andrieu C, Jeanny JC, Gollomb G and Behar- 
Cohen FF (2005) Nanoparticles for gene delivery. Mol Vis 17:124-132. 
79. El-Samaligy M, Ronjanasakul Y, Charlton JF, Weinstei GW and Lim JK (1996) Ocular 
disposition of nanoencapsulated acyclovir and ganciclovir. Drug Del 3:93-97. 
80. Kozak Y, Andrieux K, Villarroya H, Klein C, Thillaye-Goldenberg B, Naud MC, Garcia E 
and Couvreur P (2004) Intraocular injection of tamoxifen-loaded nanoparticles. Eur } 
Immunol 34:3702-3712. 
81. Merodio M, Irache JM, Valamanesh F and Mirshahi M (2002) Ocular disposition and 
tolerance of ganciclovir-loaded albumin nanoparticles after intravitreal injection in rats. 
Biomaterials 23:1587-1594. 
This page is intentionally left blank
28 
Nanoparticles and Microparticles as 
Vaccine Adjuvants 
Janet R. Wendorf, Manmohan Singh and 
Derek T. O'Hagan 
1. Introduction 
One of the most important current issues in vaccinology is the need for new adjuvants 
and delivery systems. Many of the vaccines currently in development are 
based on purified subunit proteins, recombinant molecules, synthetic peptides or 
plasmid DNA. Unfortunately, it is clear that this new generation of vaccines will be 
less immunogenic than traditional vaccines, and will require better adjuvants and 
delivery systems to induce optimal immune responses.1,2 In addition, non-living 
vaccines have generally proven ineffective at inducing potent cell mediated immunity 
(CMI), particularly of the Thl type. T helper cells can be classified into Th2 
and Thl subtypes, mainly based on their cytokine production profile, with Thl 
responses being characterized by the production of y interferon. Thl responses are 
likely to allow the development of vaccines against important infectious diseases, 
including HCV and HIV. 
Immunological adjuvants were originally described by Ramon3 as "substances 
used in combination with a specific antigen that produced a more robust immune 
response than the antigen alone." This broad definition encompasses a very wide 
range of materials.4 However, despite extensive evaluation of a large number of candidates 
over many years, the only adjuvants currently approved by the US Food 
and Drug Administration are aluminum based mineral salts, generically known 
675 
676 Wendorf, Singh & O'Hagan 
as alum. Alum has a good safety record but comparative studies show that it is a 
weak adjuvant for antibody induction to recombinant protein vaccines and induces 
a Th2, rather than a Thl response.5 In addition, Alum is not effective in inducing 
mucosal IgA antibody responses. Moreover, alum adjuvants can induce IgE antibody 
responses and have been associated with allergic reactions in some subjects.5,6 
Although Alum has been used as an adjuvant for many years, its mechanism of 
action remains poorly defined. It was originally thought to provide a "depot" effect, 
resulting in the persistence of antigen at the injection site. However, more recent 
studies involving radio-labeled antigens suggest that Alum does not establish a 
depot at the injection site.7 Recent work has indicated that Alum upregulates costimulatory 
signals on human monocytes and promotes the release of IL-4.8 Alum 
adsorption may also contribute to a reduction in toxicity for some vaccines, due 
to the adsorption of contaminating LPS.9 A key issue in adjuvant development 
is toxicity, since safety concerns have restricted the development of adjuvants ever 
since Alum was first introduced more than 50 years ago.10 Many experimental adjuvants 
have advanced to clinical trials and some have demonstrated high potency, 
but most have proven too toxic for routine clinical use. For standard prophylactic 
immunization in healthy individuals, only adjuvants that induce minimal adverse 
effects will prove acceptable. Additional practical issues that are important for adjuvant 
development include biodegradability, stability, ease of manufacture, cost and 
applicability to a wide range of vaccines. 
Adjuvants can be used to improve the immune response to vaccine antigens 
in several different ways, including (1) increasing the immunogenicity of weak 
antigens, (2) enhancing the speed and duration of the immune response, (3) modulating 
antibody avidity, specificity, isotype or subclass distribution, (4) stimulating 
CTL, (5) promoting the induction of mucosal immunity, (6) enhancing immune 
responses in immunologically immature, or senescent individuals, (7) decreasing 
the dose of antigen in the vaccine to reduce costs, or (8) helping to overcome antigen 
competition in combination vaccines. 
The mechanisms of most adjuvants still remains poorly understood, since 
immunization often activates a complex cascade of responses, and the principle 
mechanism of the adjuvant is often difficult to discern clearly. However, if one 
accepts the geographical concept of immune reactivity, in which antigens that do 
not reach the local lymph nodes do not induce responses,11 it becomes easier to 
propose mechanistic interpretations for some adjuvants, particularly those based 
on a "delivery" mechanism such as nanoparticles and microparticles. If antigens 
which do not reach lymph nodes do not induce responses, then any adjuvant which 
enhances delivery of antigen into the cells that traffic to the lymph node may 
enhance the response. A subset of dendritic cells (DC) are thought to be the key 
Nanoparticles and Microparticles as Vaccine Adjuvants 677 
cells which circulate in peripheral tissues and act as "sentinels", being responsible 
for the uptake of antigens and their transfer to lymph nodes, where they are then 
presented to T cells. Circulating immature DC are efficient for antigen uptake, while 
mature DC are efficient at antigen presentation to T cells. Hence, promoting antigen 
uptake into DC, trafficking to lymph nodes and DC maturation are thought to be 
the key components to the generation of potent immune responses. DC are thought 
to be the most effective antigen presenting cells (APC), although macrophages can 
also function in this role. 
It can be argued that the role of adjuvants for recombinant vaccines is to 
ensure that the vaccine resembles infection closely enough to initiate a potent 
immune response.12 In addition, the innate immune system directs the balance 
of humoral and CMI,13 and adjuvants can control the type of acquired immune 
response induced.14 Adjuvants can be divided into different broad groups based 
on their principal modes of action, depending on whether or not they have direct 
immunostimulatory effects on APC, or whether they function as antigen delivery 
systems. Particulate adjuvants (e.g. emulsions, microparticles, iscoms, liposomes, 
virosomes and virus-like particles) have comparable dimensions to the pathogens 
which the immune system evolved to combat. Immunostimulatory adjuvants may 
also be included in particulate delivery systems to enhance the level of response, or 
to focus the response through a desired pathway, e.g. Thl. In addition, formulating 
potent immunostimulatory adjuvants into delivery systems through restricting the 
systemic circulation of the adjuvant may limit adverse events. Nanoparticle and 
microparticle adjuvants generally act as delivery systems, although the materials 
they are made of may also have some adjuvant effect. 
In the studies from 1976 onwards, Kreuter et al}5 described the use of polymeric 
nanoparticles (50 nm to 300 nm in size) as adjuvants for adsorbed and 
entrapped vaccines. However, the polymethyl methacrylates used in these studies 
are degraded in vivo only very slowly. Faster degrading particles prepared with 
the more biocompatible poly-lactic acid (PLA) and poly (lactide-co-glycolide) (PLG) 
polymers have subsequently been extensively investigated as adjuvants. Size was 
shown to be an important parameter affecting the immunogenicity of microparticles, 
since smaller particles (< 10/xm) were significantly more immunogenic than 
larger ones.16,17 With PLG particles of size 1-10/xm (mean of 3.5/xm) compared 
with 10-110 /xm (mean of 54.5 /xm) with entrapped staphylococcal enterotoxin B, 
the serum IgG response was much higher with the smaller particles.16 With a model 
antigen entrapped in PLA particles, there was an increased antibody response with 
particles < 5 /xm compared with particles with mean sizes larger than 5 /xm.18 The 
effect of particle size on immunogenicity is likely to be a consequence of enhanced 
uptake into lymphatics and greater uptake into antigen presenting cells for the 
678 Wendorf, Singh & O'Hagan 
smaller sized particles, since only microspheres < 5 /zm were transported to the 
spleen.19 The advantages of particles with a mean size of less than 5 microns 
for optimal immune responses has been demonstrated on a number of occasions, 
but the data supporting the use of nanoparticles is less convincing. In addition to 
the inherent immunogenicity of nano- versus microparticles, carrier capacity and 
the efficiency of antigen entrapment in different formulations also needs careful 
consideration. 
The dividing line between nanoparticles and microparticles is ill defined, with 
some sources considering 1000 nm particles to be nanoparticles,20 while the United 
States patent office has the class definition for nanotechnology using the scale 
1-100 nm or slightly larger. In this chapter, the evidence supporting the advantages 
of nanoparticles versus microparticles will be critically assessed. Previous 
reviews have focused on alternative delivery routes, including nasal21 and 
oral immunization.22 In addition, the use of nanoparticles for DNA vaccines,23 
the stability of vaccines following microencapsulation,24'25 the cellular uptake of 
nanoparticles and microparticles,26 and the overall role of adjuvants in the immune 
response27,28 has also been reviewed. Although a wide variety of microparticles 
and nanoparticles have been developed as possible vaccine adjuvants, we will 
focus primarily on systems where in vivo immune responses have been reported 
following systemic immunization, although mucosal immunization will be briefly 
covered. 
2. Nanoparticle and Microparticle Preparation Methods 
There are many biodegradable or biologically compatible polymers that have been 
used for the preparation of nanoparticles and microparticles. Different methods 
have been used for particle formation and these methods have optimal size ranges 
as well as suitability for use with different polymeric agents. The preparation methods 
used, the typical size ranges, polymeric materials and antigens that have been 
evaluated are summarized in Table 1. 
2.1. Nanoparticles and microparticles made from polyesters 
Nanoparticles and microparticles made from poly (lactide-co-glycolide) (PLG) and 
poly (lactide) (PLA) and their derivatives have been widely investigated for vaccine 
delivery. Using preformed polymers, particles with entrapped antigens can be 
prepared from a variety of emulsification evaporation methods. These methods are 
based on the formation of a multiple emulsion (water in oil in water emulsion) from 
which the oil phase, an organic solvent used to dissolve the polymer, is evaporated, 
resulting in the preparation of an aqueous suspension of particles. This technique 
Table 1 Summary of various particle preparation methods using different poly 
Material 
PLG or PLA 
PLG or PLA 
poly(cyanoacrylate) 
subramolecular 
biovectors 
chitosan 
chitosan-DNA 
carbon nanotubes 
calcium phosphate 
Method 
emulsification 
evaporation 
solvent displacement 
chain polymerization 
crosslinking 
sodium sulfate 
precipitation 
complexation 
modification for 
covalent attachment 
various 
Size Range 
200nm-10/i.m 
80 nm-500 nm 
90 nm-800 nm 
60 nm 
300 nm-1 /xm 
20 nm-500 nm 
nano 
100 nm-1.2//m 
Antigen Type 
DNA, protein, 
small 
molecules 
protein, small 
molecules 
protein 
protein 
protein 
DNA 
protein, small 
molecules 
DNA, protein 
smalle 
increas 
concen 
smalle 
polym 
concer 
monom 
compl 
with co 
undete 
compa 
680 Wendorf, Singh & O'Hagan 
generally has a lower limit for the particle preparation of about 0.5 /xm (500 nm), 
although 200 nm particles can be created with a low PLG concentration and an 
increased surfactant concentration.29 PLG particles, with encapsulated model proteins 
having diameters as small as 320 nm, can also be formed by sonicating the 
emulsions.30 Various alternative approaches have also been described based on 
emulsions, including spray drying31 and phase separation.32,33 However, one of 
the drawbacks of microencapsulation of antigens is the instability of the antigen 
due to the exposure to solvents and the high shear force during microparticle 
preparation.25,34 An alternate approach to encapsulation is to adsorb the antigen 
onto the microparticles after the particle has been formed, avoiding the exposure 
of the antigen to solvents and high shear.35-37 
The solvent displacement method (also sometimes referred to as interfacial 
deposition) was first described by Fessi et a/.38,39 and allows the preparation of 
nanoparticles from preformed polymers. A water-miscible solvent (i.e. acetone) is 
used to dissolve the polymer with magnetic stirring, which is then added to an 
aqueous solution. The nanoparticles are formed by diffusion and the solvent is 
eliminated by evaporation. Depending on the solvent, polymer type, polymer concentration 
and addition of emulsifiers, these particles can range in size from 80 nm to 
500 nm or larger.40,41 The antigens for encapsulation need to be water soluble and 
compatible with the water-miscible solvent. However, the encapsulation efficiency 
is often low for water-soluble molecules.42 Nanocapsules, with an internal oil core, 
can be formed when a small volume of oil is introduced into the organic phase, and 
these can be used to dissolve less water-soluble antigens, and to increase encapsulation 
efficiency.38 However, one limitation of this approach is the preparation of low 
particle concentrations, due to the dilute initial polymer concentrations necessary 
in producing 100 nm particles. In addition to PLG, other polymers used to prepare 
nanoparticles include poly-e-caprolactone (PCL)43,44 and sulfobutylated poly(vinyl 
alcohol)-g-PLG.36,45 
Chain polymerization of modified cyanoacrylate monomers has also been used 
to make particles with a diameter of 100 nm,46 and with an additional polysaccharide 
coating, particles ranging in size from 93 to 800 nm can be produced.47 Polymerization 
preparation of nanoparticles has also been used for methyl methacrylate 
polymers.15 However, with the polymerization approach to nanoparticles preparation, 
there are concerns about the levels of residual monomers in the final formulation. 
Supramolecular Biovectors (™SMBV) are positively charged particles 
with a polysaccharide core surrounded by a phospholipidic layer, with a mean size 
of 60 nm.48,49 They are made from maltodextrins that were crosslinked with 2,3- 
epoxychloropropane and branched with glycidyl trimethylammonium to form a 
gel, which was then homogenized to give the nanoparticles, with the lipids added 
for the layer outside. Antigens were then adsorbed to the nanoparticles. 
Nanoparticles and Microparticles as Vaccine Adjuvants 681 
2.2. Nanoparticles and microparticles made with chitosan 
Chitosan is a natural biodegradable polymer of glucosamine and N-acetylglucosamine. 
It is made from the partial de-acetylation of chitin which is found in the shells 
of crustaceans.50 It has been shown to be an effective adjuvant for the intranasal 
delivery of vaccines, enhancing T-cell response and antibody levels when used 
as a soluble polymer.51 Chitosan is cationic and readily binds negatively charged 
materials, including DNA and the sialic acid found on cell surfaces. 
Chitosan and chitosan coated particles can be made using several methods. 
Using a chitosan with a low molecular weight, 700 nm chitosan particles can 
be formed by sodium sulfate precipitation. With sonication, the particles can be 
reduced in size to 300 nm.52 Chitosan coated poly-e-caprolactone particles were 
made using the solvent displacement method described above, with sizes ranging 
from 230-500 nm.44 With an emulsification method, chitosan was dissolved in the 
external phase to form 500 nm PLG particles with a chitosan coating.53 
Nanoparticles in the range of 20 nm to 500 nm can be formed spontaneously 
upon mixing of chitosan with DNA. The zeta potential (surface charge) can vary 
from negative to positive, depending on the ratio of DNA to chitosan, although 
some particles were unstable and showed aggregation.54 
2.3. Other nanoparticles and microparticles 
Functionalized carbon nanotubes have been investigated as particles for vaccine 
delivery. Through organic modification, multiple sites for covalent attachment can 
be made available for small molecules, sugars, peptides or proteins.55 The biological 
compatibility of carbon nanotubes is not certain yet, since they are not biodegradable. 
Another nanoparticle preparation method uses emulsifying wax of cetyl alcohol 
and polysorbate in combination with SDS to create microemulsions, which are 
then cooled to form particles ranging in size for 90 to 425 nm, depending on the 
SDS concentration.56 Calcium phosphate particles have been studied for vaccine 
delivery, but are usually > 10 /zm in size. However, calcium phosphate nanoparticles 
have also been studied, although the reported mean size was < 1.2 /xm, and 
these formulations were called "nanoparticles".57 Alternative calcium phosphate 
particles of mean size 100-120 nm have been used to encapsulate plasmid DNA for 
gene therapy application.58 
3. Adjuvant Effect of Nanoparticles and Microparticles 
The adjuvant effect achieved as a consequence of the association of antigens with 
particles has been known for many years.15 The enhanced immunogenicity of 
particulate antigens is unsurprising, since pathogens are particulates of similar 
682 Wendorf, Singh & O'Hagan 
dimensions and the immune system has evolved to deal with these.59 Particulate 
delivery systems present multiple copies of antigens to the immune system and 
promote trapping and retention of antigens in the local lymph nodes. Moreover, 
particles are taken up by macrophages and dendritic cells, leading to enhanced 
antigen presentation and the release of cytokines, so as to promote the induction 
of an immune response. Antigen uptake by APC is enhanced by the association of 
antigen with polymeric particles, or by the use of polymers or proteins which selfassemble 
into particles. A particularly attractive feature of particles is their ability to 
control the rate of release of entrapped antigens. Many alternative antigen delivery 
systems that are available are particulates, including liposomes, ISCOM's, micelles 
and emulsions.59,60 
Aluminum adjuvants have several limitations which has encouraged the search 
for alternative approaches. Aluminium adjuvants are not effective for all antigens, 
induce some local reactions, induce IgE antibody responses and generally 
fail to induce cell-mediated immunity, particularly cytotoxic T-cell responses. In 
the early studies, microparticles with entrapped protein61,62 and peptide63 antigens 
were shown to induce cytotoxic T lymphocyte (CTL) responses in mice following 
systemic61 and mucosal immunization.62 Microparticles also induced a delayedtype 
hypersensitivity (DTH) response,62 which is thought to be mediated by Thl 
cells, and potent T cell proliferative responses. The limited data available on the 
induction of cytokine responses in cells from animals immunized with microencapsulated 
antigens indicates that microparticles preferentially induce a Thl type 
response.61,64 Hence, microparticles may possess some inherent advantages over 
the more established Alum based adjuvants. Macrophages have been reported to be 
responsible for phagocytosis and the presentation of particulate antigens through 
the cytosolic MHC class I restricted pathway.65 However, dendritic cells are also 
likely to play an important role in the presentation of particulate antigens and the 
release of cytokines to promote a Thl type response.66 
The effect of particle size on antibody induction and cell mediated immunity 
has been investigated, and it has been concluded that 1 /xm particles are generally 
better than larger ones. However, the data supporting the benefit of nanoparticles 
over microparticles is considerably less convincing. 
3.1. Nanoparticles and microparticles as mucosal adjuvants 
Mucosal administration of vaccines offers a number of advantages over the traditional 
approach to vaccine delivery, which normally involves systemic injection 
using a needle and syringe. Mucosal delivery would eliminate the possibility of 
infections caused by inadequately sterilized needles or the re-use of needles. Also, 
mucosal vaccines might result in the induction of mucosal immunity at the sites 
Nanoparticles and Microparticles as Vaccine Adjuvants 683 
where many pathogens initially infect hosts. Mucosal delivery most commonly 
involves oral and intranasal (i.n.) immunization, although alternate routes are also 
available. The potency of particles for mucosal delivery is generally dependent on 
their ability to be taken up across the mucosal epithelium. In many studies, the 
uptake of particles by the mucosal associated lymphoid tissues (MALTs) of the 
Peyer's patches in the gut and the MALT of the respiratory tract have been demonstrated, 
albeit a very inefficient process. Moreover, there is good evidence that the 
composition of the particles impacts efficiency of uptake, including evidence that 
the binding of PLG microparticles to M cells of the Peyer's patches is less efficient 
than the binding of latex particles.67,68 
A number of alternative approaches have been evaluated for the mucosal delivery 
of vaccines using particulate carriers of various mean sizes. Chitosan particles of 
300-350 nm with associated tetanus toxoid (TT) were administered i.n. in mice, and 
induced significantly higher serum IgG responses compared with free antigen.69 In 
addition, sulfobutylated poly (vinyl alcohol)-g-PLG particles of mean size 100 nm 
with adsorbed TT were administered orally and i.n., as well as enhanced serum IgA 
(and IgG for the i.n.) antibodies, compared with soluble controls.36 SMBV nanoparticles 
with Hepatitis B surface antigen administered i.n. induced cytotoxic T lymphocyte 
(CTL) responses and higher serum IgG antibody responses than soluble 
protein.70 Calcium phosphate particles (size < 1.2 /xm) were used for the mucosal 
delivery of a herpes simplex virus type 2 antigen and they induced greater mucosal 
IgA and IgG, and systemic IgG responses, compared with soluble antigen.57 However, 
the calcium phosphate particles were sized before the final protein coating, 
so it is unclear what size the particles were when they were actually administered. 
Overall, although these various observations support the contention that particulate 
antigens are better than soluble antigens for mucosal delivery, all approaches 
appear to fall short of any likelihood of commercial development. Moreover, the 
rationale for preparing nanoparticles rather than the more established microparticles 
is not necessarily clear. 
Data from studies evaluating the effect of particle size on the induction of 
mucosal immunity, following mucosal administration of vaccines, has offered 
conflicting outcomes, depending upon the specific polymeric system evaluated 
(Table 2). In one study, with the model antigen ovalbumin (OVA) adsorbed to chitosan 
particles of varying sizes (400 nm, 1 /xm, 3 /an) for i.n. administration, higher 
IgA responses were seen with 400 nm and 1 /u,m particles, compared with the 3 /xm 
particles. However, there was no difference in the response with the 400 nm and the 
1 /x,m sized particles.52 The antigen adsorption efficiency was similar for all particles, 
although no information is reported on the antigen release profile. The conclusion 
that smaller nanoparticles were more immunogenic than larger microparticles was 
not confirmed by the paper. 
684 Wendorf, Singh & O'Hagan 
Table 2 Summary of various particles with different sizes showing the mucosal adjuvant 
effect. 
Particle 
Material 
chitosan 
PLG 
PLG 
PEG-PLA 
sulfobutylated 
pva-g-plg 
sulfobutylated 
pva-g-plg 
chitosan 
SMBV 
calcium 
phosphate 
Antigen 
model 
protein 
(ovalbumin) 
model 
protein 
(BSA) 
model 
protein 
(BSA) 
protein 
(TT) 
protein 
(TT) 
protein 
(TT) 
protein 
(TT) 
protein 
(HBsAg) 
protein 
(HSV-2) 
Sizes 
400 nm, 
1 /xm, 
3/xm 
200 nm, 
500 nm, 
1000 nm 
200 nm, 
500 nm, 
1000 nm 
200 nm, 
1.5 ixvtx 
100 nm, 
500 nm, 
1500 nm 
100 nm, 
500 nm, 
1500 nm 
350 nm 
60 nm 
<1.2/Ltm 
Route 
i.n. 
i.n. 
oral 
i.n. 
oral 
i.n. 
i.n. 
i.n. 
mucosal 
Result 
comparable IgA for 
smaller sizes, both 
greater than 
control / 3 ^im 
IgG responses of 
1000 nm> 500 nm > 
200 nm 
IgG responses of 
1000 n m > 500 n m ~ 
200 nm 
comparable IgG and 
IgA responses for 
both sizes 
IgG and IgA 
responses of 
100nm>500nm, 
none for 1500 nm 
comparable IgG/IgA 
response of 100 nm, 
500 nm, none for 
1500 nm 
higher IgG and IgA 
compared to free 
antigen 
high CTL and IgG 
compared to free 
protein 
higher IgG and IgA 
compared to free 
antigen 
Reference 
[52] 
[71] 
[71] 
[72] 
[36] 
[36] 
[69] 
[70] 
[57] 
It is also known that particle charge is important for particles to be transferred 
to the APCs and the uptake by the MALT.68 This would suggest that positively 
charged chitosan particles may behave differently than negatively charged particles. 
Therefore, similar responses with micron and sub-micron particles may not 
apply to all particle types. Chitosan particles are biodegradable and may have 
an inherent immunostimulatory effect which pther polyesters (PLGA) lack. Thus, 
Chitosan may be a promising candidate for a particulate system of size 1 micron, 
although there is no justification for chitosan nanoparticles at this moment. 
Another study evaluated the effect of particle size, using a specialized method 
of size 100 nm with adsorbed TT administered orally and i.n. This nanoparticle 
Nanoparticles and Microparticles as Vaccine Adjuvants 685 
formulation induced increased serum IgA and IgG antibody responses in comparison 
with soluble antigen control.36 Kamm et al.36 also examined 100, 500 and 
1500 nm particles. Following oral administration, the highest serum IgG and IgA 
responses were found with 100 nm particles and the responses were progressively 
lower for the 500 nm and 1500 nm nanoparticles. Following nasal administration, 
the 100 nm and 500 nm particles induced comparable IgG and IgA responses, which 
were higher than the responses with 1500 nm particles. The authors speculated that 
the different size dependence observed for the different routes is due to the different 
translocation mechanisms in the NALT, as compared with the GALT. The 
overall observations were that sub-micron nanoparticles were more immunogenic 
than larger particles. The SBPVA-PLG may be preferred to PLA for its faster degradation 
rate, however, the grafting chemistry required, renders it much less suitable 
for commercial development. 
In contrast to the studies described above, it has also been claimed in some 
studies that 1 /im particles are more effective than nanoparticles. With BSA encapsulated 
into biodegradable "nanospheres" administered i.n. and orally, 1000 nm 
PLG particles elicited higher serum IgG responses than 200 nm and 500 nm particles 
(71). For i.n. administration, the 500 nm particles also induced higher serum 
IgG responses than the 200 nm particles. The IgG2a/IgGl ratios with the different 
particle sizes were similar and higher than antigen alone and that with alum.71 From 
in vitro release studies, it was found that the 1000 nm particles did have a different 
release profile from the 200 nm and 500 nm particles, which may account for some 
of the differences seen in vivo. For the oral studies, the authors hypothesized that 
200 nm and 500 nm particles are more readily absorbed through the intestinal wall 
than 1000 nm particles, absorbed almost exclusively by Peyer's patch cells, leading 
to higher immune responses for the larger particles. The two studies comparing 
particle size,36,71 reached differing conclusions, but it is unclear whether this was 
due to the differences in polymers used for the studies, the different antigens used, 
or that of different formulations. 
Particle type as well as particle size can also have a strong influence on immune 
responses. Using 200 nm PEG-PLA particles, 200 nm PLA particles and 1.5 /xm PEGPLA 
particles with encapsulated TT administered i.n., it was found that the 200 nm 
and 1.5 /an PEG-PLA induced higher IgG and IgA antibody responses, compared 
with the PLA particles.72 Although the two particle sizes were comparable, but the 
PEG-PLA particles performed better than PLA particles of the same size. The PEGPLA 
particles are less hydrophobic than the PLA particles. This may explain some 
of the differences as particle uptake is strongly influenced by the hydrophobicity of 
the polymer. The authors also hypothesized that the difference between the particle 
types is related to the propensity of the PLA nanoparticles to aggregate in vitro, 
indicative of the relative stabilities of the particles in the mucosal fluids. Although 
686 Wendorf, Singh & O'Hagan 
these particles are interesting, the PEG-PLA polymer is not commercialy available, 
thereby hindering development. The material of the particle system is an important 
factor to consider when comparing particles. 
3.2. Nanoparticles and microparticles as systemic adjuvants 
The use of nanoparticles as systemic adjuvants is summarized in Table 3. The advantage 
of particulates for vaccine delivery, compared with soluble antigens or alum, 
has been investigated in a number of studies and has been extensively evaluated 
by several groups (2,24). In one study, TT and CpG (a known immunostimulatory 
oligonucleotide) were co-encapsulated within PLG nanospheres of mean size 290- 
310 nm and were evaluated in mice. This formulation resulted in the induction of an 
enhanced antigen specific T-cell proliferative response, in comparison with the soluble 
antigen plus CpG.73 The co-delivery of TT and CpG within PLG nanoparticles 
induced very strong serum IgG response. This is another instance of nanoparticles 
out performing soluble antigen. However, the rationale for preparing nanoparticles 
rather than the more established microparticles is not necessarily clear. 
Table 3 Summary of various particles with different sizes showing the systemic adjuvant 
effect of these formulations. 
Particle 
Material 
Antigen Sizes Route Result Reference 
chitosan 
PLG 
PLA 
carbon 
nanotubes 
PLG 
model protein 
(ovalbumin) 
model protein 
(BSA) 
protein (TT) 
peptide (fmdv) 
protein (TT) + 
CpG 
400 nm, 
1 [im 
200 nm, 
500 nm, 
1000 nm 
630 nm, 
4 Aim 
diameter 
~15-60nm 
length 
~500nm 
290-310 nm 
s.c. 
..p. 
s.c. 
comparable IgG [52] 
for both sizes and 
greater than 
control 
IgG responses of [71] 
1000nm> 
500nm~200nm 
comparable titers, [76] 
both sizes less 
effective than 
alum 
more neutralizing [55] 
antibodies 
compared to free 
antigen 
higher IgG, [73] 
Thltype 
compared to free 
antigen; 
enhanced T-cell 
proliferation 
Nanoparticles and Microparticles as Vaccine Adjuvants 687 
Similarly, in another study, using an alternate carbon polymer, a nonbiodegradable, 
inorganic nano-scale particle was investigated. Carbon nanotubes 
with covalently linked peptides from a foot-and-mouth disease virus were administered 
to mice i.p. and the carbon nanotubes elicited high levels of virus-neutralizing 
antibodies, compared with the free peptide control.55 Carbon nanotubes introduce 
many problems. They are not biodegradable and may be toxic; hence they are not 
a good choice for vaccine formulation.74,75 
Data from studies evaluating the effect of particle size on the induction of systemic 
immunity have offered conflicting outcomes. A single dose of PL A particles 
with TT encapsulated (4/xm and 630nm) induced lower anti-TT titers than two 
doses of alum. There were no significant difference between the larger and smaller 
particles.76 The lower response with the particle formulation may be due to the 
limitation of having only one dose in particles, as compared with two doses on 
the alum. The lower response with the PLA formulation may also be due to the 
encapsulation process compromising of antigen stability.34,77 
The effect of particle size using BS A as an antigen was also evaluated by Gutierro 
et al.n The model antigen BSA was encapsulated into PLG particles of varying 
sizes (i.e. lOOOnm, 200 nm and 500 nm) and administered s.c. This elicited higher 
serum IgG with the lOOOnm, compared with the smaller particles.71 The in vitro 
release profile was different for lOOOnm PLG particles which may account for the 
differences in responses. In another similar study, no size dependent effect was 
found with a model antigen OVA adsorbed to chitosan particles of varying sizes 
400 nm, 1 /tm. Higher serum IgG responses were seen with the particles compared 
with the soluble antigen, but there was no difference between the two sizes.52 
Both studies described above used model antigens and the size conclusions 
may not be applicable beyond model antigens. Also, for some vaccines, antibody 
response may not directly correlate with protective efficacy of the vaccine. One 
major difference between these two studies is the net surface charge of the particles. 
It has been shown in vitro that positively charged particles (poly-L-lysine 
coated polystyrene) induced higher phagocytosis in dendritic cells. Surface charge 
as well as particle size might influence uptake by macrophages and dendritic cells.78 
Therefore, the difference in size dependent behavior observed may also be more 
prominent involving particles of a specific charge. 
Recent studies have also shown that nanoparticles may exert an adjuvant effect 
for the induction of cell mediated immunity compared with soluble antigen.70,79 
Particle size may be an important parameter influencing the efficacy of microparticles 
as adjuvants for CTL induction. Nixon et al.63 showed that microparticles 
< 500 nm induce better CTL responses than microparticles > 2 /xm.63 In another 
study, a non-degradable polystyrene nanoparticle with OVA antigen covalently 
bound was administered i.d. in mice. It was seen that 40 nm particles induced 
688 Wendorf, Singh & O'Hagan 
highest T-cell responses in comparison with the 2 /xm size.80 However, in this paper, 
the initial particle size is reported, but the size after covalent attachment of the antigen 
is not reported. Therefore, the conclusion cannot be confirmed by the findings 
reported. 
4. Delivery of DNA Using Nanoparticles and Microparticles 
The previous sections were mainly focused on the delivery of proteins or peptides, 
but DNA nanoparticles and microparticles have also been used. Nanoparticle studies 
with DNA as adjuvants are summarized in Table 4. DNA plasmids are weakly 
immunogenic and particles may help boost the immune response. Cationic PLG 
particles with adsorbed CTAB of sizes 300 nm, 1 /tm, and 30 /xm were used to adsorb 
an HIV-1 DNA plasmid and delivered i.m. The 300 nm and 1 /xm particles induced, 
significantly enhanced IgG titers, compared with naked DNA and the 30 /xm particles. 
The 1 /xm particles were capable of inducing potent CTL responses, whereas 
naked DNA failed to induce CTL activity81 It appeared in this study that the 300 nm 
particles were better than the 1 /xm particles. An additional study with more animals 
confirmed that the 1 /xm and 300 nm PLG particles with adsorbed plasmid 
DNA, induced comparable IgG serum titers (Fig. 1). There was no advantage for 
the smaller sized particles. It is believed that the efficient delivery of DNA to APCs is 
Table 4 Various particle based delivery systems for DNA showing the mucosal systemic 
adjuvant effect of these formulations. 
Particle 
Material 
PLG (CTAB) 
chitosan 
chitosan 
polylysine-gimidazoleatic 
acid 
Antigen 
DNA (HIV-1) 
DNA (peanut 
allergen) 
DNA 
(tuberculosis) 
DNA (HIV-2 
env) 
Sizes 
300 nm, 
1 /xm, 3 /xm 
150-300 nm 
376 nm 
140 nm 
Route 
i.m. 
oral 
pulmonary 
i.d. 
Result 
higher IgG for 
300 nm, 1 /xm 
compared to free 
DNA 
higher IgA and 
IgG2a response 
compared to free 
DNA 
increased IFN-y 
compared to 
DNA alone of i.m. 
Administration 
higher IgM, IgG, 
IgA compared to 
free antigen 
Reference 
[81] 
[82] 
[83] 
[84] 
Nanoparticles and Microparticles as Vaccine Adjuvants 689 
105 - 
104 - 
o 103H 
icr 
101 - 
10° 
J 300 nm PLG DNA 
I7~71 1 nm PLG DNA 
L~ 3 soluble DNA 
_J_ 
_I_ 
00 nm 
1 H9 
1 |im 
1 H9 
soluble 
1 l-iQ 
300 nm 
10 ng 
1 |i,m 
10 ng 
soluble 
10 ng 
Fig. 1. Serum IgG titers in groups (w = 20 for particles, n = 10 for soluble) of BALB/c mice 
immunized with either PLG-CTAB-p55 gag DNA of size 300 nm or 1 /xm or DNA alone at 
two dose levels of 10 /xg and 1 /xg. Antibody titers are geometric mean titers ± SE at 2 weeks 
post-second immunization (week 6) time point after immunizations at day 0 and day 28. 
The response from the 300 nm and 1 /xm particles at both dose levels are not significantly 
different. 
an important component of the adjuvant effect, since larger microparticles > 30 /xm 
did not elicit a strong immune response. 
The association of DNA with nanoparticles and microparticles has been shown 
to be more effective than naked DNA. Chitosan-DNA particles (150-300 nm) of 
peanut allergen gene delivered orally, produced secretory IgA and serum IgG2a, 
compared with no detectable response with naked DNA.82 The pulmonary delivery 
of chitosan-DNA particles (average size of 376 nm) with plasmid DNA encoding 
T-cell epitopes from mycobacterium tuberculosis, were able to induce the maturation 
of dendritic cells and the increased level of IFN-y secretion, compared with 
the plasmid in solution or i.m. delivery.83 In another DNA study, polylysine-graft 
imidazoleacetic acid complexed with DNA with a diameter of 140 nm was used 
for the HIV env plasmid.84 It was found that IgG, IgM and IgA responses were 
increased several folds, compared with naked DNA. It was also speculated that 
this formulation may also help protect the DNA from nuclease degradation.84 Based 
690 Wendorf, Singh & O'Hagan 
on the evidence in the literature, the rationale for preparing nanoparticles rather 
than the more established microparticles is not necessarily clear with the DNA 
formulations. 
5. Conclusions 
A number of systems with different types of antigen (proteins, peptides, DNA) 
have been investigated with the particles ranging in size from 50 nm to 1000 nm. 
For most systems, the critical particle size is < 5 /xm, with particles in the range 
of 100 nm to 1 /xm, often inducing comparable immune responses. In some cases, 
depending on the route of delivery, there may be increased immune responses with 
the smaller nanoparticles. For i.n. administration, there was no evidence that 100 nm 
particles were better than micron particles. For oral administration, some studies 
found enhanced responses with nanoparticles, compared with microparticles, while 
other studies found equivalence, or that microparticles elicited higher responses. 
Overall, the evidence for nanoparticles (~100nm) outperforming microparticles 
(1-2 /xm) for enhanced immunogenicity is weak. Further examination is needed to 
support nanoparticles as a better formulation in place of microparticles. Also, some 
of the studies carried out were done model antigens and the same results may/may 
not apply to relevant antigens where vaccine efficacy is determined. However, there 
are some advantages to nanoparticles compared with microparticles that have not 
been directly addressed. For instance, smaller nanoparticles (sub-200 nm) can be 
sterile filtered, allowing the particle preparation to be a non-sterile process with 
terminal sterilization. 
The distinction between nanoparticles and microparticles is usually made by 
the authors and there is no consistency in what constitutes a nanoparticle, and 
this needs careful consideration when comparing results from the literature. The 
important size measurement point is immediately prior to administration, postlyophilization, 
or other processing, and this is not always reported. More relevant 
endpoints such as protective efficacy may be crucial in distinguishing between 
nanoparticles and microparticles. Nanoparticles and microparticles both constitute 
a very effective vaccine delivery system. Some of these formulations are currently 
in pre-clinical and clinical evaluations. 
Acknowledgments 
We would like to acknowledge the contributions of our colleagues in Chiron Corporation 
to the ideas contained in the chapter, particularly, all the members of the 
Vaccine Adjuvants and Delivery Group. 
Nanoparticles and Microparticles as Vaccine Adjuvants 691 
References 
1. Gupta RK and Siber GR (1995) Adjuvants for human vaccines-current status, problems 
and future prospects. Vaccine 13:1263-1276. 
2. Gupta RK, Griffin P, Jr., Chang AC, Rivera R, Anderson R, Rost B, Cecchini D, 
Nicholson M and Siber GR (1996) The role of adjuvants and delivery systems in modulation 
of immune response to vaccines. Adv Exp Med Biol 397:105-113. 
3. Ramon G (1924) Sur la toxine et surranatoxine diphtheriques. Ann Inst Pasteur 38. 
4. Vogel FR and Powell MF (1995) A compendium of vaccine adjuvants and excipients. 
Pharm Biotechnol 6:141-228. 
5. Gupta RK (1998) Aluminum compounds as vaccine adjuvants. Adv Drug Deliv Rev 
32:155-172. 
6. Relyveld EH, Bizzini B and Gupta RK (1998) Rational approaches to reduce adverse 
reactions in man to vaccines containing tetanus and diphtheria toxoids. Vaccine 16: 
1016-1023. 
7. Gupta RK, Chang AC, Griffin P, Rivera R and Siber GR (1996) In vivo distribution of 
radioactivity in mice after injection of biodegradable polymer microspheres containing 
14C-labeled tetanus toxoid. Vaccine 14:1412-1416. 
8. Ulanova M, Tarkowski A, Hahn-Zoric M and Hanson LA (2001) The Common vaccine 
adjuvant aluminum hydroxide up-regulates accessory properties of human monocytes 
via an interleukin-4-dependent mechanism. Infect Immun 69:1151-1159. 
9. Shi Y, HogenEsch H, Regnier FE and Hem SL (2001) Detoxification of endotoxin by 
aluminum hydroxide adjuvant. Vaccine 19:1747-1752. 
10. Edelman R (1997) New Generation Vaccines. Adjuvants for the future, in Levine GCW, 
MM, Kaper JB and Cobon GS (eds.). Marcel Dekker, Inc., New York, pp. 173-192. 
11. Zinkernagel RM, Ehl S, Aichele P, Oehen S, Kundig T and Hengartner H (1997) Antigen 
localization regulates immune responses in a dose- and time-dependent fashion: A 
geographical view of immune reactivity. Immunol Rev 156:199-209. 
12. Fearon DT (1997) Seeking wisdom in innate immunity. Nature 388:323-324. 
13. Fearon DT and Locksley RM (1996) The instructive role of innate immunity in the 
acquired immune response. Science 272:50-53. 
14. Yip HC, Karulin AY, Tary-Lehmann M, Hesse MD, Radeke H, Heeger PS, Trezza RP, 
Heinzel FP, Forsthuber T and Lehmann PV (1999) Adjuvant-guided type-1 and type-2 
immunity: Infectious/noninfectious dichotomy defines the class of response. / Immunol 
162:3942-3949. 
15. Kreuter J and Speiser PP (1976) New adjuvants on a polymethylmethacrylate base. Infect 
Immun 13:204-210. 
16. Eldridge JH, Staas JK, Meulbroek JA, Tice TR and Gilley RM (1991) Biodegradable and 
biocompatible poly (DL-lactide-co-glycolide) microspheres as an adjuvant for staphylococcal 
enterotoxin B toxoid which enhances the level of toxin-neutralizing antibodies. 
Infect Immun 59:2978-2986. 
17. O'Hagan DT, Jeffery H and Davis SS (1993) Long-term antibody responses in mice 
following subcutaneous immunization with ovalbumin entrapped in biodegradable 
microparticles. Vaccine 11:965-969. 
692 Wendorf, Singh & O'Hagan 
18. Nakaoka R, Inoue Y, Tabata Y and Ikada Y (1996) Size effect on the antibody production 
induced by biodegradable microspheres containing antigen. Vaccine 14:1251-1256. 
19. Tabata Y, Inoue Y and Ikada Y (1996) Size effect on systemic and mucosal immune 
responses induced by oral administration of biodegradable microspheres. Vaccine 
14:1677-1685. 
20. Quintanar-Guerrero D, Allemann E, Fessi H and Doelker E (1998) Preparation techniques 
and mechanisms of formation of biodegradable nanoparticles from preformed polymers. 
Drug Dev Ind Pharm 24:1113-1128. 
21. Vajdy M and O'Hagan DT (2001) Microparticles for intranasal immunization. Adv Drug 
Del Rev 51:127-141. 
22. Singh M and O'Hagan D (1998) The preparation and characterization of polymeric antigen 
delivery systems for oral administration. Adv Drug Del Rev 34:285-304. 
23. Cui Z and Mumper RJ (2003) Microparticles and nanoparticles as delivery systems for 
DNA vaccines. Crit Rev Ther Drug Can Syst 20:103-137. 
24. Gupta RK, Singh M and O'Hagan DT (1998) Poly(lactide-co-glycolide) microparticles 
for the development of single-dose controlled-release vaccines. Adv Drug Del Rev 32: 
225-246. 
25. Tamber H, Johansen P, Merkle HP and Gander B (2005) Formulation aspects of 
biodegradable polymeric microspheres for antigen delivery. Adv Drug Del Rev 57: 
357-376. 
26. Panyam J and Labhasetwar V (2003) Biodegradable nanoparticles for drug and gene 
delivery to cells and tissue. Adv Drug Del Rev 55:329-347. 
27. Degen WG, Jansen T and Schijns VE (2003) Vaccine adjuvant technology: From mechanistic 
concepts to practical applications. Expert Rev Vaccines 2:327-335. 
28. Schijns VE (2000) Immunological concepts of vaccine adjuvant activity: Commentary. 
Cur Opin Immun 12:456-463. 
29. Scholes PD, Coombes AGA, Ilium L, Daviz SS, Vert M and Davies MC (1993) The preparation 
of sub-200 ran poly(lactide-co-glycolide) microspheres for site-specific drug delivery. 
/ Control Rel 25:145-153. 
30. Blanco MD and Alonso MJ (1997) Development and characterization of protein-loaded 
poly(lactide-co-glycolide) nanospheres. Eur } Pharm Biopharm 43:287-294. 
31. Nguyen XC, Herberger JD and Burke PA (2004) Protein powders for encapsulation: A 
comparison of spray-freeze drying and spray drying of darbepoetin alfa. Pharm Res 
21:507-514. 
32. Thomasin C, Merkle HP and Gander B (1998) Drug microencapsulation by PLA/PLGA 
coacervation in the light of thermodynamics. 2. Parameters determining microsphere 
formation. / Pharm Sci 87:269-275. 
33. Thomasin C, Ho NT, Merkle HP and Gander B (1998) Drug microencapsulation by 
PLA/PLGA coacervation in the light of thermodynamics. 1. Overview and theoretical 
considerations. / Pharm Sci 87:259-268. 
34. Cleland JL and Jones AJ (1996) Stable formulations of recombinant human growth hormone 
and interferon-gamma for microencapsulation in biodegradable microspheres. 
Pharm Res 13:1464-1475. 
Nanoparticles and Microparticles as Vaccine Adjuvants 693 
35. Singh M, Kazzaz J, Chesko J, Soenawan E, Ugozzoli M, Giuliani M, Pizza M, 
Rappouli R and O'Hagan DT (2004) Anionic microparticles are a potent delivery system 
for recombinant antigens from Neisseria meningitidis serotype B. J Pharm Sci 93:273-282. 
36. Jung T, Kamm W, Breitenbach A, Hungerer KD, Hundt E and Kissel T (2001) Tetanus 
toxoid loaded nanoparticles from sulfobutylated polyvinyl alcohol)-graft-poly(lactideco-
glycolide): Evaluation of antibody response after oral and nasal application in mice. 
Pharm Res 18:352-360. 
37. Kazzaz J, Neidleman J, Singh M, Ott G and O'Hagan DT (2000) Novel anionic microparticles 
are a potent adjuvant for the induction of cytotoxic T lymphocytes against recombinant 
p55 gag from HIV-1. / Control Rel 67:347-356. 
38. Fessi H, Puisieux F, Devissaguet JP, Ammoury N and Benita S (1989) Nanocapsule formation 
by interfacial polymer deposition following solvent displacement, lnt } Pharm 
55:R1-R4. 
39. Fessi H, Puisieux F and Devissaguet JP (1987) Process for the preparation of dispersible 
colloidal systems of a substance in the form of nanocapsules. Eur Patent. 
40. Peltonen L, Koistinen P, Karjalainen M, Hakkinen A and Hirvonen J (2002) The 
effect of cosolvents on the formulation of nanoparticles from low-molecular-weight 
poly(l)lactide. AAPS Pharm Sci Tech 3:E32. 
41. Wehrle P, Magenheim B and Benita S (1995) The influence of process parameters on the 
pla nanoparticle size distribution, evaluated by means of factorial design. Eur } Pharm 
Biopharm 41:19-26. 
42. Niwa T, Takeuchi H, Hino T, Kunou N and Kawashima Y (1993) Preparations of 
biodegradable nanospheres of water-soluble and insoluble drugs with — lactide/ 
glycolide copolymer by a novel spontaneous emulsification solvent diffusion method, 
and the drug release behavior. / Control Rel 25:89-98. 
43. Molpeceres J, Guzman M, Aberturas MR, Chacon M and Berges L (1996) Application 
of central composite designs to the preparation of polycaprolactone nanoparticles by 
solvent displacement. / Pharm Sci 85:206-213. 
44. Calvo P, RemunanLopez C, Vilajato JL and Alonso MJ (1997) Development of positively 
charged colloidal drug carriers: Chitosan coated polyester nanocapsules and submicronemulsions. 
Coll Polym Sci 275:46-53. 
45. Jung T, Breitenbach A and Kissel T (2000) Sulfobutylated poly(vinyl alcohol)-graftpoly(
lactide-co-glycolide)s facilitate the preparation of small negatively charged 
biodegradable nanospheres. / Control Rel 67:157-169. 
46. O'Hagan DT, Palin KJ and Davis SS (1989) Poly(butyl-2-cyanoacrylate) particles as adjuvants 
for oral immunization. Vaccine 7:213-216. 
47. Chauvierre C, Labarre D, Couvreur P and Vauthier C (2003) Novel polysaccharidedecorated 
poly(isobutyl cyanoacrylate) nanoparticles. Pharm Res 20:1786-1793. 
48. Major M, Prieur E, Tocanne JF, Betbeder D and Sautereau AM (1997) Characterization 
and phase behaviour of phospholipid bilayers adsorbed on spherical polysaccharidic 
nanoparticles. Biochim Biophys Acta 1327:32^0. 
49. Baudner BC, Balland O, Giuliani MM, Von Hoegen P, Rappuoli R, Betbeder D and Del 
Giudice G (2002) Enhancement of protective efficacy following intranasal immunization 
694 Wendorf, Singh & O'Hagan 
with vaccine plus a nontoxic LTK63 mutant delivered with nanoparticles. Infect Immun 
70:4785-4790. 
50. Singla AK and Chawla M (2001) Chitosan: Some pharmaceutical and biological aspectsan 
update. / Pharm Pharmacol 53:1047-1067. 
51. McNeela EA, Jabbal-Gill I, Ilium L, Pizza M, Rappuoli R, Podda A, Lewis DJM 
and Mills KHG (2004) Intranasal immunization with genetically detoxified diphtheria 
toxin induces T cell responses in humans: Enhancement of Th2 responses and toxinneutralizing 
antibodies by formulation with chitosan. Vaccine 22:909-914. 
52. Nagamoto T, Hattori Y, Takayama K and Maitani Y (2004) Novel chitosan particles and 
chitosan-coated emulsions inducing immune response via intranasal vaccine delivery. 
Pharm Res 21:671-674. 
53. Vila A, Sanchez A, Tobio M, Calvo P and Alonso MJ (2002) Design of biodegradable 
particles for protein delivery. / Control Rel 78:15-24. 
54. Ilium L, Jabbal-Gill I, Hinchcliffe M, Fisher AN and Davis SS (2001) Chitosan as a novel 
nasal delivery system for vaccines. Adv Drug Del Rev 51:81-96. 
55. Pantarotto D, Partidos CD, Hoebeke J, Brown F, Kramer E, Briand J-P, Muller S, Prato 
M and Bianco A (2003) Immunization with Peptide-Functionalized Carbon Nanotubes 
Enhances Virus-Specific Neutralizing Antibody Responses. Chem Biol 10:961-966. 
56. Cui Z and Mumper RJ (2002) Coating of cationized protein on engineered nanoparticles 
results in enhanced immune responses. Int J Pharm 238:229-239. 
57. He Q, Mitchell A, Morcol T and Bell SJ (2002) Calcium phosphate nanoparticles induce 
mucosal immunity and protection against herpes simplex virus type 2. Clin Diagn Lab 
Immunol 9:1021-1024. 
58. Bisht S, Bhakta G, Mitra S and Maitra A (2005) pDNAloaded calcium phosphate nanoparticles: 
Highly efficient non-viral vector for gene delivery. Int} Pharm 288:157-168. 
59. O'Hagan DT (1994) Novel Delivery Systems for Oral Vaccines. Microparticles as oral vaccines. 
O'Hagan DT, Editor. CRC Press, Boca Raton, 175. 
60. O'Hagan DT, Ott GS and Van Nest G (1997) Recent advances in vaccine adjuvants: The 
development of MF59 emulsion and polymeric microparticles. Mol Med Today 3:69-75. 
61. Moore A, McGuirk P, Adams S, Jones WC, McGee JP, O'Hagan DT and Mills KH (1995) 
Immunization with a soluble recombinant HIV protein entrapped in biodegradable 
microparticles induces HIV-specific CD8+ cytotoxic T lymphocytes and CD4+ Thl cells. 
Vaccine 13:1741-1749. 
62. Maloy KJ, Donachie AM, O'Hagan DT and Mowat AM (1994) Induction of mucosal 
and systemic immune responses by immunization with ovalbumin entrapped in 
poly(lactide-co-glycolide) microparticles. Immunology 81:661-667. 
63. Nixon DF, Hioe C, Chen PD, Bian Z, Kuebler P, Li ML, Qiu H, Li XM, Singh M, 
Richardson J, McGee P, Zamb T, Koff W, Wang CY and O'Hagan D (1996) Synthetic 
peptides entrapped in microparticles can elicit cytotoxic T cell activity. Vaccine 14: 
1523-1530. 
64. Vordermeier HM, Coombes AG, Jenkins P, McGee JP, O'Hagan DT, Davis SS and Singh 
M (1995) Synthetic delivery system for tuberculosis vaccines: Immunological evaluation 
of the M. tuberculosis 38 kDa protein entrapped in biodegradable PLG microparticles. 
Vaccine 13:1576-1582. 
Nanoparticles and Microparticles as Vaccine Adjuvants 695 
65. Kovacsovics-Bankowski M and Rock KL (1995) A phagosome-to-cytosol pathway for 
exogenous antigens presented on MHC class I molecules. Science 267:243-246. 
66. Scheicher C, Mehlig M, Dienes HP and Reske K (1995) Uptake of microparticle-adsorbed 
protein antigen by bone marrow-derived dendritic cells results in up-regulation of 
interleukin-1 alpha and interleukin-12 p40/p35 and triggers prolonged, efficient antigen 
presentation. Eur ] Immunol 25:1566-1572. 
67. Brayden DJ and Baird AW (2001) Microparticle vaccine approaches to stimulate mucosal 
immunisation. Microbes Infect 3:867-876. 
68. Florence AT (1997) The oral absorption of micro- and nanoparticulates: Neither exceptional 
nor unusual. Phtzrm Res 14:259-266. 
69. Vila A, Sanchez A, Janes K, Behrens I, Kissel T, Jato JLV and Alonso MJ (2004) Low 
molecular weight chitosan nanoparticles as new carriers for nasal vaccine delivery in 
mice. EurJPharm Biopharm 57:123-131. 
70. Debin A, Kravtzoff R, Santiago JV, Cazales L, Sperandio S, Melber K, Janowicz Z, 
Betbeder D and Moynier M (2002) Intranasal immunization with recombinant antigens 
associated with new cationic particles induces strong mucosal as well as systemic 
antibody and CTL responses. Vaccine 20:2752-2763. 
71. Gutierro I, Hernandez RM, Igartua M, Gascon AR and Pedraz JL (2002) Size dependent 
immune response after subcutaneous, oral and intranasal administration of BSA loaded 
nanospheres. Vaccine 21:67-77. 
72. Vila A, Sanchez A, Evora C, Soriano I, Jato JLV and Alonso MJ (2004) PEG-PLA nanoparticles 
as carriers for nasal vaccine delivery. / Aerosol Med 17:174-185. 
73. Diwan M, Tafaghodi M and Samuel J (2002) Enhancement of immune responses 
by co-delivery of a CpG oligodeoxynucleotide and tetanus toxoid in biodegradable 
nanospheres. / Control Rel 85:247-262. 
74. Lam C-W, James JT, McCluskey R and Hunter RL (2004) Pulmonary Toxicity of Single- 
Wall Carbon Nanotubes in Mice 7 and 90 Days After Intratracheal Instillation. Toxicol 
Sci. 77:126-134. 
75. Jia G, Wang HF, Yan L, Wang X, Pei RJ, Yan T, Zhao YL and Guo XB (2005) Cytotoxicity 
of carbon nanomaterials: Single-wall nanotube, multi-wall nanotube, and fullerene. 
Environ Sci Tech 39:1378-1383. 
76. Katare YK, Panda AK, Lalwani K, Haque IU and Ali MM (2003) Potentiation of immune 
response from polymer-entrapped antigen: Toward development of single dose tetanus 
toxoid vaccine. Drug Del 10:231-238. 
77. Johnson OL, Cleland JL, Lee HJ, Charnis M, Duenas E, Jaworowicz W, Shepard D, 
Shahzamani A, Jones AJ and Putney SD (1996) A month-long effect from a single injection 
of microencapsulated human growth hormone. Nat Med 2:795-799. 
78. Thiele L, Rothen-Rutishauser B, Jilek S, Wunderli-AUenspach H, Merkle HP and Walter 
E (2001) Evaluation of particle uptake in human blood monocyte-derived cells in vitro. 
Does phagocytosis activity of dendritic cells measure up with macrophages? / Control 
Rel 76:59-71. 
79. Cui Z, Patel J, Tuzova M, Ray P, Phillips R, Woodward JG, Nath A and Mumper RJ (2004) 
Strong T cell type-1 immune responses to HIV-1 Tat (1-72) protein-coated nanoparticles. 
Vaccine 22:2631-2640. 
696 Wendorf, Singh & O'Hagan 
80. Fifis T, Gamvrellis A, Crimeen-Irwin B, Pietersz GA, Li J, Mottram PL, McKenzie IF 
and Plebanski M (2004) Size-dependent immunogenicity: Therapeutic and protective 
properties of nano-vaccines against tumors. / Immunol 173:3148-3154. 
81. Singh M, Briones M, Ott G and O'Hagan D (2000) Cationic microparticles: A potent 
delivery system for DNA vaccines. Proc Natl Acad Sci USA 97:811-816. 
82. Roy K, Mao HQ, Huang SK and Leong KW (1999) Oral gene delivery with 
chitosan — DNA nanoparticles generates immunologic protection in a murine model of 
peanut allergy. Nat Med 5:387-391. 
83. Bivas-Benita M, van Meijgaarden KE, Franken KLMC, Junginger HE, Borchard G, Ottenhoff 
THM and Geluk A (2004) Pulmonary delivery of chitosan-DNA nanoparticles 
enhances the immunogenicity of a DNA vaccine encoding HLA-A*0201-restricted Tcell 
epitopes of Mycobacterium tuberculosis. Vaccine 22:1609-1615. 
84. Locher CP, Putnam D, Langer R, Witt SA, Ashlock BM and Levy JA (2003) Enhancement 
of a human immunodeficiency virus env DNA vaccine using a novel polycationic 
nanoparticle formulation. Immunol Lett 90:67-70. 
29 
Pharmaceutical Nanocarriers in 
Treatment and Imaging of Infection 
Raymond M. Schiffelers, Cert Storm and 
Irma A. J. M. Bakker-Woudenberg 
Nanoscaled carrier systems can be used in the treatment and imaging of infectious 
diseases. To optimize accumulation of the carrier at the site of infection, the characteristics 
of the nanocarriers should complement the pathophysiological processes 
that play a role during infection. 
As carriers are recognized as foreign materials, they can be employed for targeting 
the mononuclear phagocyte system (MPS). The MPS consists of cells that are 
specialized in the clearance of the body of foreign particles. As such, these cells are 
involved in the clearance of microorganisms and they form an important replication 
site for infectious organisms to survive intracellularly. 
If carriers can avoid recognition by the MPS, it allows them to take advantage 
of the enhanced capillary permeability that is one of the hallmarks of the inflammatory 
reaction coinciding with infection. Together with the local efflux of plasma, 
nanoscaled carriers can enter the inflamed area through convective forces. 
Finally, the carrier can be locally administered to interact specifically with tissues, 
cells or microorganisms that are present at the site of infection. 
1. Introduction 
Pharmaceutical nanocarrier systems include liposomes, nanoparticles, micelles 
and emulsions, and they can be used in the treatment and imaging of infectious 
diseases.1-3 These systems are used to improve the degree of localization or the 
697 
698 Schiffelers, Storm & Bakker-Woudenberg 
persistence of encapsulated antimicrobial drugs or imaging molecules at the site 
of infection by altering the molecules' pharmacokinetics and tissue distribution 
profiles. To optimize accumulation of the carrier at the site of infection, factors on 
the side of the carrier, as well as on the side of the infected host should be taken 
into account. Ideally, the characteristics of the nanocarriers should be tailored to 
complement the pathophysiological processes that play a role during infection.4-6 
Three strategies to target sites of infection have been employed: (1) the use of 
carriers that are recognized as foreign materials; (2) the use of carriers that avoid 
recognition as foreign materials, and (3) local delivery of carriers. These targeting 
strategies are discussed in the light of recent developments in delivery strategies 
and the introduction of new antimicrobial agents such as the nubiotics for bacterial 
infections and siRNA for viral infections. 
2. Carriers that are Easily Recognized as Foreign Materials 
Most carriers that are administered in vivo are almost always rapidly recognized as 
foreign materials. Proteins in biological fluids such as activated complement components, 
opsonize the carriers to facilitate recognition and uptake7-9 by cells that 
constitute the MPS. These cells are continuously surveiling the body to detect and 
phagocytose these opsonized foreign particles. This avid recognition and uptake 
mechanism can be employed for targeting purposes. These cells are of interest since 
many infectious agents are also recognized as foreign materials and end up in the 
same type of cells. Microorganisms that are able to survive intracellularly may be 
difficult to reach for conventional antimicrobial treatment, due to the barriers posed 
by the cellular membranes.10-13 
One of the first demonstrations of the value of nanoscale carriers for the delivery 
of drugs to the MPS was delivered by New and colleagues as early as 1978.14 
Liposome-encapsulated antimony was used in this study to treat experimental 
leishmaniasis, showing several orders of magnitude increase in the therapeutic 
index. Leishmaniasis is still an endemic disease in several parts of the world and also 
occurs as an opportunistic infection in immunocompromised patients. Nowadays, 
treatment of leishmaniasis is complicated by widespread resistance to antimony 
compounds leading to treatment failure and relapse. Lipid formulations of amphotericin 
B are currently employed in the clinic as possible option for the treatment of 
visceral and cutaneous leishmaniasis.15-18 
The lipid formulation is used to facilitate drug accumulation in the 
macrophages. As such, it is preferred over free amphotericin B deoxycholate 
(Fungizone®). The nanoparticulate structures in this formulation are unstable in 
the blood stream, leading to higher toxicity especially in the kidneys.16,19 
Pharmaceutical Nanocarriers in Treatment and Imaging of Infection 699 
Oil-in water-emulsions of submicron sizes, also known as lipid nanospheres 
that have been loaded with piperine, have been described as a new formulation 
for the treatment of visceral leishmaniasis.20 All nanosphere formulations applied 
were more effective than the free drug in reducing parasite counts in all organs 
investigated in a model of Leishmania donovani infection in Balb/c mice at a dose of 
5 mg/kg. The most effective formulation also contained the cationic lipid stearylamine. 
Inclusion of charged lipids usually promotes phagocyte uptake of lipidic 
carrier systems, which probably explains the increased therapeutic effect.21 Alternatively, 
stearylamine has been described as leishmanicidal agent.22 Furthermore, 
it was shown that enzyme levels in blood (as a measure for liver toxicity) and creatinine, 
and urea levels in the blood (as a measure for renal toxicity) were not altered 
after administration of the drug in the nanosphere formulations, indicating that 
toxicity of these formulations in these organs is limited. 
Recently, a liposomal pentavalent antimony formulation was described, targeted 
at the macrophage via the scavenger receptors A and B for the treatment 
of Leishmania chagasi amastigotes.23 The natural affinity of these receptors for 
polyanions was used by the inclusion of the negatively charged phospholipid phosphatidylserine 
in the liposomal membrane. The negative charge mediates receptormediated 
endocytosis. Furthermore, the authors demonstrated that the scavenger 
receptor was upregulated upon infection, possibly through secretion of transforming 
growth factor-/il, which could preferentially increase uptake by infected 
macrophages. Nevertheless, infected macrophages ingested less liposomes than 
normal macrophages, possibly because of the high parasite burden per macrophage 
that could affect normal cell metabolism. Overall, the targeted drug showed a 
16-fold higher efficacy, compared with free compound in vitro. 
The same strategy with phosphatidylserine was used to target liposomes to 
macrophages infected with Cryptococcus neoformans. Incorporation of chloroquine 
in liposomes resulted in increased delivery to macrophages in vitro and in vivo. 
This promoted infected mouse survival and reduced parasite counts in liver and 
brain.24 
Mannose was used as a recognition signal for macrophage uptake by the mannose 
receptor by Medda et al.25 Phospholipid microspheres consisting of polylacticco-
glycolic acid and phosphatidyl ethanol amine were grafted with mannose 
and subcutaneously injected. In this system, dihydroindolo [2,3-a] indolizine, an 
antileishmanial agent, was incorporated. The drug was administered at a dose 
of 3 mg/kg. In Leishmania donovani infected hamsters, mannose-grafted microspheres 
suppressed parasite numbers in the spleen over 90%, whereas free drug 
only reduced numbers by 26%. Furthermore, it was shown that infection induced 
changes in the blood levels of hepatic enzymes, creatinine and urea. These changes 
could largely be prevented by the formulated drug. 
700 Schiffelers, Storm & Bakker-Woudenberg 
Mannose-coated liposomes were also employed to deliver CpG-containing 
oligodeoxynucleotides to macrophages infected with Leishmania donovani.26 These 
CpG-oligonucleotides activate macrophages through the presence of Cytidine 
phosphorothioate Guanosine islands that are strong activators of macrophages 
via Toll-like receptors, leading to interleukin-12 and interferon-)/ production. In 
mice, suffering from visceral leishmaniasis, mannose-coated liposomes containing 
CpG-oligonucleotides, were more effective in inhibiting amastigote growth than 
liposome-encapsulated CpG-oligonucleotides or free CpG-oligonucleotides. In the 
spleen, the mannose-coated liposomal formulation of the CpG-oligonucleotides 
completely eliminated the parasites, whereas both controls failed to achieve complete 
eradication. 
A dual targeting approach combining tuftsin and nystatin in liposomes was 
reported by Khan et al.27,28 In their studies, the effects of the liposomal antifungal 
agents on Candida albicans infection in mice was studied. The tuftsin-bearing 
liposomes were taken up by macrophages by binding to the tuftsin receptor which 
activated these cells. As these cells were loaded with liposomal nystatin, this may 
improve their killing of phagocytosed C. albicans organisms through the added 
action of the drug. Alternatively, the macrophages may function as a depot for subsequent 
prolonged release of the drug. In this study, combination of tuftsin and 
nystatin in the liposome formulation increased C. albicans clearance in liver, spleen, 
kidneys and lungs, and reduced the numbers of organisms in the blood stream, 
thereby promoting mice survival, compared with free agents or liposomal formulation 
of either tuftsin or nystatin alone. 
In another study, delivery of immunomodulatory compounds to macrophages 
was demonstrated. The important role of macrophages in generating proinflammatory 
reactions had been the basis for targeting antisense molecules to 
these cells.29 Antisense oligomers against tumor necrosis factor-alpha were incorporated 
into albumin microspheres. Endotoxin was administered i.v. or Escherichia 
coli was administered i.p. to rats. Animals received 100 fig antisense TNF and TNFinhibition 
and rat survival was measured. Microencapsulated antisense improved 
survival in these endotoxin and E. coli peritonitis models, compared with free antisense. 
Importantly, delivery by albumin encapsulation prolonged the action of the 
antisense to 72 hrs. 
In a similar approach, Sioud and Sorensen employed the process of RNA interference 
to silence TNF-production.30 The process was based on cellular delivery of 
small interfering RNA (siRNA), small double-stranded RNA molecules that mediate 
degradation of complementary mRNA sequences. siRNA was administered at a 
dose of 50 to 100 /xg and was complexed to liposomes that were based on the cationic 
lipid dioleoyl trimethylammonium propane. The complexes were injected i.p. and 
shown to prevent the induction of TNF production upon challenge by endotoxin. 
Pharmaceutical Nanocarriers in Treatment and Imaging of Infection 701 
The specificity of the inhibition was demonstrated by the fact that the induction of 
IL-6 expression remained unchanged. 
3. Carriers that Avoid Recognition as Foreign Materials 
The specific accumulation of carriers that are easily recognized as foreign materials 
by macrophage can offer great benefits when the infectious organism is exclusively 
localized in these cells. For micro-organisms that are primarily present in other 
tissues, this fast recognition can become a major obstacle to reach distant sites. 
Therefore, carriers have been developed to avoid recognition by the MPS, allowing 
these carrier systems to take advantage of the increased capillary permeability at 
the site of infection.31-34 The increased capillary permeability as well as vascular 
leakage are accompanying inflammatory reactions together with infection. With the 
local efflux of plasma, plasma proteins and immune cells, nanoscaled carriers can 
enter the inflamed area through convective forces. This approach has been used in 
the treatment and imaging of infections. 
In magnetic resonance based imaging, the natural tropism of > 150 nm iron 
oxide particles to be rapidly taken up by liver and spleen macrophages prevented 
imaging of peripheral infections. In the studies by Kaim et al., the usage of ultrasmall 
nanoparticles was explored.35,36 The nanoparticles consisted of a core of iron 
oxide crystals that is coated with dextran. The mean particle size was 35 nm. As 
a result of this small particle size, the blood half-life in rats increased to approximately 
2-3 hrs. The authors demonstrated successful imaging of soft tissue infections 
in rats using this system. They postulate that the mechanism responsible is 
the uptake by peripheral macrophages in the blood stream, followed by extravasation 
at the site of infection based on the absence of non-phagocytosed particles 
in the inflammatory milieu. However, the presence of non-ingested particles may 
be underestimated due to the cellular accumulation, making it difficult to quantify 
individual extracellular particles. 
Long-circulating liposomes, which have a substantially prolonged circulation 
time compared with the iron oxide particles that were described above (a half-life 
of approximately 24 hrs in rats), have also been successfully employed to image 
a variety of infectious diseases. The use of long-circulating liposomes for imaging 
purposes has been reviewed recently.37-39 
Recent advances include the detection of invasive pulmonary aspergillosis at 
an early stage of infection.40 The causative fungus Aspergillus fumigatus is difficult to 
detect using conventional imaging agents, and in most cases, only at an advanced 
stage of the disease. Furthermore, interpretation of CT-scan images is difficult as 
the supposedly characteristic halo or air crescent-sign of the infection is not always 
clear. Imaging of intravenously injected 99mTc-PEG- liposomes in a rat model of 
702 Schiffelers, Storm & Bakker-Woudenberg 
left-sided invasive pulmonary aspergillosis, demonstrated that 82% of the scintigraphic 
images revealed the presence of the fungus already at 48-hr inoculation. 
Active infection was needed for the accumulation of labeled liposomes to occur, as 
inoculation with saline or killed Aspergillus-spores did not cause increased liposome 
accumulation in the inoculated lung. 
The long circulatory half-life of PEG-liposomes results in a gradual increase 
in accumulation at the site of infection, which is in general considered an advantage 
as the scintigraphic image improves. Nevertheless, in specific situations when 
the infection is localized in the vicinity of the heart or large blood vessels, the 
background activity of the blood remains high for prolonged periods of time, thus 
hampering visualization. Laverman et al. demonstrated that biotin-coated 99mTc- 
PEG-liposomes kept their long-circulating half-life.41 However, upon injection of 
avidin, complexes were formed intravenously, resulting in a rapid clearance from 
the blood stream by MPS-organs. This clearance coincided with a loss of activity 
in the blood pool, allowing visualization of an experimental abscess in the neck 
region of the rabbits, previously undetected due to the activity in the heart region. 
These long-circulating PEG-liposomes can also be loaded with antimicrobial 
agents to achieve site-specific drug delivery, as has been recently reviewed.31'42,43 
In our own laboratory, we have evaluated the factors on the side of the host, as 
well as on the side of the liposome that determine target localization. It appeared 
that the area under the blood concentration time-curve determined the degree of 
localization at the site of infection; whereas the level of vascular leakage was the 
most important determinant on the side of the host.44-47 Optimized PEG-liposomes 
were loaded with gentamicin and therapeutic effects were studied in a rat Klebsiella 
pneumoniae model. Rat survival was strongly dependent on the gentamicin formulation. 
In immunocompetent rats, liposomal gentamicin was superior over the free 
drug, even though complete therapeutic efficacy could be achieved with multiple 
administrations of the free drug. In leukopenic rats, a combination of free and 
liposomal gentamicin showed best therapeutic effects, compared with either treatment 
alone. It is postulated that liposomal gentamicin produces low therapeutically 
active drug concentrations in the blood stream, which are insufficient to control the 
rapidly occurring bacteremia in the case of impaired host defense. However, liposomal 
gentamicin does localize in the infected lung and leads to local bacterial killing. 
Therefore, the combination of free drug (producing active antimicrobial levels in 
the circulation) and liposomal drug (leading to high drug levels in the lung) showed 
optimal results.48 
PEG-liposomes were also loaded with a combination of antimicrobial agents 
that was predicted to interact synergistically based on in vitro data.49 Whether 
synergistic drug action in vivo also occurs when the agents are administered 
in the free form is questionable, as their differing pharmacokinetic profile and 
tissue distribution do not correlate with the static drug concentrations present 
Pharmaceutical Nanocarriers in Treatment and Imaging of Infection 703 
in the in vitro assays. However, the use of a targeted drug carrier such as longcirculating 
PEG-liposomes, could enforce a parallel tissue distribution resulting 
in increased concentrations at the site of infection. Gentamicin and ceftazidime 
were co-encapsulated in long-circulating liposomes, as these agents were demonstrated 
in vitro to act synergistically on K. pneumoniae. These liposomes were tested 
for their ability to prolong the survival of rats infected in the left lung with 
a high gentamicin/ceftazidime-susceptible or a gentamicin/ceftazidime-resistant 
K. pneumoniae. The high susceptible K. pneumoniae could be effectively treated 
with single doses of lipsomal-gentamicin or liposomal-ceftazidime. Liposomal coencapsulation 
of both agents allowed reductions in doses due to a synergistic 
therapeutic effect of these antibiotics. In the resistant K. pneumoniae model, the 
co-encapsulated agents again resulted in the synergistic activity of both antibiotics, 
allowing effective treatment of the otherwise lethal pneumonia. Interestingly, in 
both models, combinations of the two free drugs were merely additive. 
O-stearylamylopectin was used as a recognition signal for the uptake of PEGliposomes 
by lung macrophages.50 The macrophages in this organ are an important 
replication site for Mycobacterium tuberculosis replication. In a murine model 
of M. tuberculosis infection, administration of liposome-encapsulated isoniazid and 
rifampicin reduced the number of bacteria in lungs, liver and spleen more efficiently 
than the free drugs. There was no significant reduction in bacterial numbers 
in mice treated with free drugs once weekly for 6 weeks, compared with untreated 
control mice. Liposomal treatment resulted in approximately 90% decrease in bacterial 
numbers. The enhanced efficacy of the liposome-encapsulated drugs is likely 
to enhance delivery to macrophages in liver, spleen, and lungs. 
In another study focused on osteomyelitis, vancomycin and ciprofloxacin were 
encapsulated in liposomes.51 As compared with neutral or anionic liposomes, 
cationic liposomes entrapped the highest percentage of drugs and showed highest 
antibacterial activity in vitro, probably due to the charge based cell interactions that 
can occur in vitro. The cationic liposomes were tested for therapeutic efficacy in 
a rabbit model of chronic staphylococcal osteomyelitis. Both free ciprofloxacin or 
vancomycin over a period of two weeks failed to achieve bone sterilization after 
i.v. injection. Combination of free ciprofloxacin and vancomycin was more effective 
only at the expense of renal dysfunction and severe diarrhea. Complete sterilization 
of the bone was seen in the group treated for two weeks with the combination 
of drugs in liposomal form, while nephrotoxicity and diarrhea were less frequent. 
Although the cationic liposomes showed preferable characteristics in vitro regarding 
drug encapsulation and antibacterial effect, it is likely that the use of neutral 
or PEG-coated liposomes would have improved results as the localization at the 
site of infection for cationic liposomes is expected to be limited due to the rapid 
clearance of charged liposome species. 
704 Schiffelers, Storm & Bakker-Woudenberg 
Another application of long-circulating liposomes is that they can be employed 
as a sustained drug delivery system in the blood stream. When ciprofloxacin is 
encapsulated in the PEG-liposomes, there is a gradual release of antibiotic from 
the liposomes.52,53 As such, the liposomal formulation displayed a prolonged presence 
in the blood and tissues of ciprofloxacin. As the encapsulated ciprofloxacin 
is shielded from the tissues by the liposomal membrane, the antibiotic could also 
be administered at relatively high doses. Interestingly, in the rat pneumonia model 
with high ciprofloxacin susceptible K. pneumoniae, 90% animal survival could not 
be achieved with the free drug at a once daily dosing schedule where the highest 
doses could be administered. In contrast, liposome-encapsulated ciprofloxacin 
injected once daily was still effective at this low frequency dose schedule. 
To promote localization of amphotericin B at infectious foci, a number of formulations 
have been devised trying to achieve prolonged presence of amphotericin 
B in the blood stream and to shield the toxicity of the drug. Espuelas and colleagues 
used poly(epsilon-caprolactone) nanospheres coated with poloxamer 188 or mixed 
micelles with of this surfactant to deliver amphotericin B.54 Both formulations had 
an approximately 10-fold lower activity in vitro against C. albicans; however, the 
activity towards infected macrophages was similar to that of Fungizone® that is 
5-fold reduced. However, the reduced toxicity was paralleled by a 4-fold reduced 
efficacy. Similar observations have been noted in most studies concerning alternative 
formulations of amphotericin B, achievement of the reduction of the toxicity of 
the drug parallels a reduction in therapeutic efficacy. 
A study by Fukui et al., however, shows that these two phenomena can be 
uncoupled.55 Lipid nanospheres were used to deliver amphotericin B, and this new 
formulation was compared with the commercially available lipid formulations of 
amphotericin B. Plasma amphotericin B-levels of Amphocil® or Abelcet® were low, 
reaching levels below 1 /xg/ml within minutes after intravenous injection in rats at 
a dose of 1 mg/kg. Amphotericin B levels in the lipid nanosphere formulation were 
nearly two orders of magnitude higher and similar to those yielded by AmBisome®. 
Interestingly, in dogs, plasma amphotericin B concentrations after administration 
in the nanosphere formulation were approximately 3-fold higher than that of 
AmBisome®. In a rat model of local candidiasis, amphotericin B-containing 
nanospheres significantly inhibited the growth of C. albicans, whereas AmBisome® 
did not, even though local amphotericin B concentrations were similar. These results 
were similar as obtained in vitro where nanopshere-incorporated amphotericin B 
was as effective as Fungizone®, while AmBisome® activity was reduced. 
Similarly, amphotericin B-containing nanoparticles were designed coated with 
heparin to achieve localization at the infected site, at the site of lesions.56 The 
heparin-coated formulation did achieve 3-fold higher concentrations in the lungs of 
mice with pulmonary blastomycosis than Fungizone®. However, this could reflect 
a difference in pharmacokinetics rather than the specific binding at the target site. In 
Pharmaceutical Nanocarriers in Treatment and Imaging of Infection 705 
this mouse model, Fungizone® dosed at the maximum tolerated dose (1.2mg/kg) 
failed to cure mice, whereas the nanoparticles dosed at 4.8mg/kg achieved a 50% 
cure rate. 
4. Local Application of Carriers 
Nanoparticulate drug delivery systems may also be applied locally at the site of 
infection to increase drug delivery or to act as a local depot from which the drug is 
gradually released. 
A new class of DNA or RNA-based antimicrobial agents (known as "nubiotics") 
was investigated by Dale et al.57 The compounds are thought to act antimicrobially 
by virtue of their proton donor capacity. The hydrogen ions would then be 
responsible for the killing of the bacteria due to membrane depolarization. However, 
the exact mechanism of action remains yet to be confirmed. In a burn-wound 
infection caused by Pseudomonas aeruginosa-model, the therapeutic effects of neutral 
liposome-encapsulated nubiotics were investigated. It appeared that intravenously 
and subcutaneously injected liposome-encapsulated drug at a dose of 20mg/kg 
was able to promote the survival of mice, whereas free nubiotics or PBS failed. 
Nucleic acid based therapeutics were also employed to treat respiratory syncytial 
virus.58 The nucleic acid was an siRNA that was intracellularly produced in 
the host by an encoding DNA-vector. The siRNA was directed towards the viral 
NS1 gene. Cell entry of this siRNA will prevent the protein from being synthesized, 
hence inhibiting viral replication. The DNA-vector was complexed to oligomeric 
nanometer-size chitosan particles and administered intranasally. siRNA against 
NS1 were shown to be produced in the lung tissues and protected against respiratory 
syncytial virus multiplication for at least 4 days. These studies show that this 
formulation may have prophylactic potential to prevent viral infection. 
The results obtained, however, should be interpreted critically, as polylCLC 
and CpG oligonucleotides are able to mount antiviral responses that are not specific; 
and the phenomena of the resulting interferon production is also noted in the 
study using siRNA described above.59'60 Prophylactic administration of liposomeencapsulated 
polylCLC completely protected mice against a lethal respiratory challenge 
of influenza A virus in mice, whereas all animals died in the control group.61 
The antiviral effect was shown to last up to 3 weeks after administration. The strong 
aspecific activity of liposome-loaded nucleic acids cautions against the straightforward 
attribution of therapeutic effects to specific mechanisms. 
A series of studies used local application of nanoparticles to address Helicobacter 
pylori infection. One study employed mucoadhesive gliadin nanoparticles, 
containing amoxicillin for increased retention in the stomach, for the eradication 
of H. pylori.62 Rhodamine labeled particles were administered to rats by gavage 
to test their gastric mucoadhesive properties. It was found that the mucoadhesive 
706 Schiffelers, Storm & Bakker-Woudenberg 
characteristics of nanoparticles increased with increasing gliadin content. In Mongolian 
gerbils, the eradication of H. pylori was evaluated after oral administration of 
amoxicillin-loaded nanoparticles. Amoxicillin alone was also able to kill H. pylori, 
but the dose needed was higher than that of the nanoparticles, likely owing to the 
prolonged presence of the particles in the stomach. 
Improvement of drug residence time was also the objective of a study using 
floating microspheres.63 Polycarbonate based particles were tested for their floating 
capacity and they were loaded with acetaminohydroxamic acid. In vivo studies on 
H. pylori-intected Mongolian gerbils showed that both free drug and drug-loaded 
particles displayed antibacterial activity in vivo, but particle encapsulation lowered 
the drug dose required for H. pylori eradication, by one order of magnitude, likely 
as a result of the prolongation of the gastric residence time of the particles. 
In another study, lipobeads were used.64 These beads consist of a polymeric particle 
core that is coated with a lipid bilayer consisting of phosphatidylethanolamine. 
The concept aims at binding the lipobeads to H. pylori phosphatidylethanolamine 
receptors. In addition, the particle is loaded with acetohydroxamic acid for gradual 
drug release in the bacteria's vicinity. In in vitro studies, the drug loaded particles 
were shown to be most efficacious in inhibiting bacterial growth and stomach cell 
adherence, compared with free drug and empty lipobeads. 
Wong et al. studied the effects of aerosolized non-PEGylated liposomes encapsulating 
ciprofloxacin.65 Levels of drug in the lungs were higher after aerosolization 
than that of free drug. The therapeutic efficacy of liposome-encapsulated 
ciprofloxacin was compared with free drug in a mouse model of pulmonary infection 
by Francisella tularensis. At 48hrs after the mice were infected intranasally 
with ten times the LD50, they were treated with aerosolized liposome-encapsulated 
ciprofloxacin or ciprofloxacin in the free form, resulting in 100% survival and 0% 
survival respectively at the doses chosen. 
Tobramycin was encapsulated in fluid liposomes that were administered 
intratracheally for the treatment of Pseudomonas aeruginosa in rats.66 Similarly, as 
described above, drug exposure in the lungs was improved upon administration of 
tobramycin in the liposomal form. Single doses of free drug or liposomal drug were 
hardly effective, as lung counts of bacteria remained >105 in 90% of the animals 
for both formulations. Only after repeated administrations, bacterial numbers < 103 
were noted in 10% of the animals treated with free drug and 30% of the animals 
treated with the liposome-encapsulated drug. 
5. Concluding Remarks 
For MPS-targeted carriers, the new developments in this category indicate an 
increasing focus on specific targeting of macrophages using receptor-mediated 
Pharmaceutical Nanocarriers in Treatment and Imaging of Infection 707 
endocytosis, possibly to increase the cell-type specificity of targeting. It may 
be expected that future studies will try to address different subpopulations of 
macrophages (infected vs non-infected, activated vs quiescent) to increase specificity 
even more. In addition, increased understanding of the immunologically 
important role of macrophages has generated interest in modulating their function 
during inflammation. Especially, the potent and highly specific new technique 
of RNA interference may offer powerful ways to modulate macrophage function. 
Nevertheless, other mechanisms such as the activation of macrophages through 
CpG-oligonucleotides may be less specific, but they can still raise a strong response 
that may have important clinical benefits. 
For long-circulating carrier systems, the gradual accumulation at the sites of 
infection due to locally increased capillary permeability is an important asset for 
the imaging studies. The use of biotin labeled liposomes to obliterate this longcirculating 
characteristic at chosen time-points by the injection of avidin is to remove 
activity from the blood pool. In certain cases, this may improve the contrast. Longcirculating 
characteristics may also be used for drug delivery at peripheral sites. 
This may also be of value for several of the microorganisms surviving inside phagocytes. 
Although liver and spleen are the prime target organs for phagocyte drug 
delivery, tissue macrophages may also be of interest where these infections are concerned. 
As long-circulating carriers are ultimately taken up by deep seated tissue 
macrophages (in addition to MPS), these carriers can also be used to reach intracellular 
infections outside the liver and spleen. 
Local application of drug delivery systems can have important therapeutic 
benefits for localized infections. The confinement of H. pylori to the stomach offers 
an ideal opportunity for local delivery. The approaches discussed in this chapter are 
aimed at increasing stomach residence time, so as to increase drug exposure of the 
pathogen. The same objective is true for the local delivery in the lung. However, 
pathogens causing pneumonia may be expected to spread from the primary site 
of infection to the distant tissues more easily, requiring additional treatment. The 
local delivery approach may also be of value for nucleic acid-based drugs that 
are otherwise, due to their inherent instability and charged character, difficult to 
deliver. 
References 
1. Allen TM and Cullis PR (2004) Drug delivery systems: Entering the mainstream. Science 
303(5665):1818-1822. 
2. Kubik T, Bogunia-Kubik K and Sugisaka M (2005) Nanotechnology on duty in medical 
applications. Curr Pharm Biotechnol 6(l):17-33. 
3. Emerich DF (2005) Nanomedicine — Prospective therapeutic and diagnostic applications. 
Exp Opin Biol Ther 5(l):l-5. 
708 Schiffelers, Storm & Bakker-Woudenberg 
4. Smith AL (2002) Inhaled antibiotic therapy: What drug? What dose? What regimen? 
What formulation? / Cyst Fibros KSuppl 2):189-193. 
5. Nissim A, Gofur Y, Vessillier S, Adams G and Chernajovsky Y (2004) Methods for targeting 
biologicals to specific disease sites. Trends Mol Med 10(6):269-274. 
6. Conway SP, Brownlee KG, Denton M and Peckham DG (2003) Antibiotic treatment of 
multidrug-resistant organisms in cystic fibrosis. Am J Respir Med 2(4):321-332. 
7. Patel HM (1992) Serum opsonins and liposomes: Their interaction and opsonophagocytosis. 
Crit Rev Ther Drug Can Syst 9(l):39-90. 
8. Szebeni J (1998) The interaction of liposomes with the complement system. Crit Rev Ther 
Drug Carrier Syst 15(l):57-88. 
9. Bradley AJ and Devine DV (1998) The complement system in liposome clearance: Can 
complement deposition be inhibited? Adv Drug Del Rev 32(l-2):19-29. 
10. van de Vosse E, Hoeve MA and Ottenhoff TH (2004) Human genetics of intracellular 
infectious diseases: Molecular and cellular immunity against mycobacteria and 
salmonellae. Lancet Inject Dis 4(12):739-749. 
11. Dussurget O, Pizarro-Cerda J and Cossart P (2004) Molecular determinants of Listeria 
monocytogenes virulence. Annu Rev Microbiol 58:587-610. 
12. Hueffer K and Galan JE (2004) Salmonella-induced macrophage death: Multiple mechanisms, 
different outcomes. Cell Microbiol 6(11):1019-1025. 
13. Soldati D, Foth BJ and Cowman AF (2004) Molecular and functional aspects of parasite 
invasion. Trends Parasitol 20(12)567-574. 
14. New RR, Chance ML, Thomas SC and Peters W (1978) Antileishmanial activity of antimonials 
entrapped in liposomes. Nature 272(5648)55-56. 
15. Goldsmith DR and Perry CM (2004) Amphotericin B lipid complex: In visceral leishmaniasis. 
Drugs 64(17):1905-1911. 
16. Veerareddy PR and Vobalaboina V (2004) Lipid-based formulations of amphotericin B. 
Drugs Today 40(2):133-145. 
17. Choi CM and Lerner EA (2002) Leishmaniasis: Recognition and management with a 
focus on the immunocompromised patient. Am J Clin Dermatol 3(2):91-105. 
18. Melby PC (2002) Recent developments in leishmaniasis. Curr Opin Infect Dis 15(5): 
485^90. 
19. Barrett JP, Vardulaki KA, Conlon C, el al. (2003) A systematic review of the antifungal 
effectiveness and tolerability of amphotericin B formulations. Clin Ther 25(5):1295-1320. 
20. Veerareddy PR, Vobalaboina V and Nahid A (2004) Formulation and evaluation of oilin-
water emulsions of piperine in visceral leishmaniasis. Pharmazie 59(3):194-197. 
21. Moghimi SM (2002) Liposome recognition by resident and newly recruited murine liver 
macrophages. / Lipos Res 12(l-2):67-70. 
22. Afrin F, Dey T, Anam K and Ali N (2001) Leishmanicidal activity of stearylamine-bearing 
liposomes in vitro.} Parasitol 87(1):188-193. 
23. Tempone AG, Perez D, Rath S, Vilarinho AL, Mortara RA and de Andrade HF, Jr (2004) 
Targeting Leishmania (L.) chagasi amastigotes through macrophage scavenger receptors: 
The use of drugs entrapped in liposomes containing phosphatidylserine. / Antimicrob 
Chemother 54(l):60-68. 
Pharmaceutical Nanocarriers in Treatment and Imaging of Infection 709 
24. Khan MA, Jabeen R, Nasti TH and Mohammad O (2005) Enhanced anticryptococcal 
activity of chloroquine in phosphatidylserine-containing liposomes in a murine model. 
J Antimicrob Chemother 55(2):223-228. 
25. Medda S, Jaisankar P, Manna RK, Pal B, Giri VS and Basu MK (2003) Phospholipid 
microspheres: Anovel delivery mode for targeting antileishmanial agent in experimental 
leishmaniasis. / Drug Targ 11(2):123-128. 
26. Datta N, Mukherjee S, Das L and Das PK (2003) Targeting of immunostimulatory DNA 
cures experimental visceral leishmaniasis through nitric oxide up-regulation and T cell 
activation. Eur J Immunol 33(6):1508-1518. 
27. Khan MA, Nasti TH, Saima K, et al. (2004) Co-administration of immunomodulator 
tuftsin and liposomised nystatin can combat less susceptible Candida albicans infection 
in temporarily neutropenic mice. FEMS Immunol Med Microbiol 41(3):249-258. 
28. Khan MA, Syed FM, Nasti HT, et al. (2003) Use of tuftsin bearing nystatin liposomes 
against an isolate of Candida albicans showing less in vivo susceptibility to amphotericin 
B. / Drug Targ ll(2):93-99. 
29. Oettinger C and D'Souza M (2003) Microencapsulation of tumor necrosis factor 
oligomers: A new approach to proinflammatory cytokine inhibition. / Interf Cytok Res 
23(9):533-543. 
30. Sioud M and Sorensen DR (2003) Cationic liposome-mediated delivery of siRNAs in 
adult mice. Biochem Biophys Res Commun 312(4):1220-1225. 
31. Bakker-Woudenberg IA (2002) Long-circulating sterically stabilized liposomes as carriers 
of agents for treatment of infection or for imaging infectious foci. Int J Antimicrob 
Agents 19(4):299-311. 
32. Woodle MC (1998) Controlling liposome blood clearance by surface-grafted polymers. 
Adv Drug Del Rev 32(1-2):139-152. 
33. Moghimi SM and Szebeni J (2003) Stealth liposomes and long circulating nanoparticles: 
Critical issues in pharmacokinetics, opsonization and protein-binding properties. Prog 
Lip Res 42(6):463^78. 
34. Ishida T, Harashima H and Kiwada H (2002) Liposome clearance. Biosci Rep 22(2): 
197-224. 
35. Kaim AH, Jundt G, Wischer T, et al. (2003) Functional-morphologic MR imaging with 
ultrasmall superparamagnetic particles of iron oxide in acute and chronic soft-tissue 
infection: Study in rats. Radiology 227(1):169-174. 
36. Kaim AH, Wischer T, O'Reilly T, et al. (2002) MR imaging with ultrasmall superparamagnetic 
iron oxide particles in experimental soft-tissue infections in rats. Radiology 
225(3):808-814. 
37. Boerman OC, Rennen H, Oyen WJ and Corstens FH (2001) Radiopharmaceuticals to 
image infection and inflammation. Semin Nucl Med 31(4):286-295. 
38. Lutz AM, Weishaupt D, Persohn E, et al. (2005) Imaging of macrophages in soft-tissue 
infection in rats: Relationship between ultrasmall superparamagnetic iron oxide dose 
and MR signal characteristics. Radiology 234(3):765-775. 
39. Ercan MT and Kostakoglu L (2000) Radiopharmaceuticals for the visualization of infectious 
and inflammatory lesions. Curr Pharm Des 6(11):1159-1177. 
710 Schiffelers, Storm & Bakker-Woudenberg 
40. Becker MJ, Dams ET, de Marie S, et al. (2002) Scintigraphic imaging using 99mTc-labeled 
PEG liposomes allows early detection of experimental invasive pulmonary aspergillosis 
in neutropenic rats. Nucl Med Biol 29(2): 177-184. 
41. Laverman P, Zalipsky S, Oyen WJ, et al. (2000) Improved imaging of infections by avidininduced 
clearance of 99mTc-biotin-PEG liposomes. / Nucl Med 41(5):912-918. 
42. Schiffelers R, Storm G and Bakker-Woudenberg I (2001) Liposome-encapsulated aminoglycosides 
in pre-clinical and clinical studies. / Antimicrob Chemother 48(3):333-344. 
43. Maurer N, Fenske DB and Cullis PR (2001) Developments in liposomal drug delivery 
systems. Exp Opin Biol Ther l(6):923-947. 
44. Schiffelers RM, Bakker-Woudenberg IA and Storm G (2000) Localization of sterically stabilized 
liposomes in experimental rat Klebsiella pneumoniae pneumonia: Dependence 
on circulation kinetics and presence of poly(ethylene)glycol coating. Biochim Biophys 
Acta 1468(l-2):253-261. 
45. Schiffelers RM, Bakker-Woudenberg IA, Snijders SV and Storm G (1999) Localization of 
sterically stabilized liposomes in Klebsiella pneumoniae-infected rat lung tissue: Influence 
of liposome characteristics. Biochim Biophys Acta 1421(2):329-339. 
46. Schiffelers RM, Storm G and Bakker-Woudenberg IA (2001) Host factors influencing 
the preferential localization of sterically stabilized liposomes in Klebsiella pneumoniaeinfected 
rat lung tissue. Pharm Res 18(6):780-787. 
47. Bakker-Woudenberg IA, Schiffelers RM, Storm G, Becker MJ and Guo L (2005) Longcirculating 
sterically stabilized liposomes in the treatment of infections. Meth Enzymol 
391:228-260. 
48. Schiffelers RM, Storm G, ten Kate MT and Bakker-Woudenberg IA (2001) Therapeutic 
efficacy of liposome-encapsulated gentamicin in rat Klebsiella pneumoniae pneumonia 
in relation to impaired host defense and low bacterial susceptibility to gentamicin. 
Antimicrob Agents Chemother 45(2):464^170. 
49. Schiffelers RM, Storm G, ten Kate MT, et al. (2001) In vivo synergistic interaction of 
liposome-coencapsulated gentamicin and ceftazidime. / Pharmacol Exp Ther 298(1): 
369-375. 
50. Labana S, Pandey R, Sharma S and Khuller GK (2002) Chemotherapeutic activity against 
murine tuberculosis of once weekly administered drugs (isoniazid and rifampicin) 
encapsulated in liposomes, hit} Antimicrob Agents 20(4):301-304. 
51. Kadry AA, Al-Suwayeh SA, Abd-Allah AR and Bayomi MA (2004) Treatment of 
experimental osteomyelitis by liposomal antibiotics. / Antimicrob Chemother 54(6): 
1103-1108. 
52. Bakker-Woudenberg IA, ten Kate MT, Guo L, Working P and Mouton JW (2002) 
Ciprofloxacin in polyethylene glycol-coated liposomes: Efficacy in rat models of acute 
or chronic Pseudomonas aeruginosa infection. Antimicrob Agents Chemother 46(8): 
2575-2581. 
53. Bakker-Woudenberg IA, ten Kate MT, Guo L, Working P and Mouton JW (2001) 
Improved efficacy of ciprofloxacin administered in polyethylene glycol-coated liposomes 
for treatment of Klebsiella pneumoniae pneumonia in rats. Antimicrob Agents 
Chemother 45(5):1487-1492. 
Pharmaceutical Nanocarriers in Treatment and Imaging of Infection 711 
54. Espuelas MS, Legrand P, Campanero MA, et al. (2003) Polymeric carriers for amphotericin 
B: In vitro activity, toxicity and therapeutic efficacy against systemic candidiasis 
in neutropenic mice. / Antimicrob Chemother 52(3):419^27. 
55. Fukui H, Koike T, Saheki A, Sonoke S and Seki J (2003) A novel delivery system for 
amphotericin B with lipid nano-sphere (LNS). Int ] Pharm 265(l-2):37-45. 
56. Clemons KV, Ranney DF and Stevens DA (2001) A novel heparin-coated hydrophilic 
preparation of amphotericin B hydrosomes. Curr Opin Investig Drugs 2(4):480^87. 
57. Dale RM, Schnell G and Wong JP (2004) Therapeutic efficacy of "nubiotics" against 
burn wound infection by pseudomonas aeruginosa. Antimicrob Agents Chemother 48(8): 
2918-2923. 
58. Zhang W, Yang H, Kong X, et al. (2005) Inhibition of respiratory syncytial virus infection 
with intranasal siRNA nanoparticles targeting the viral NS1 gene. Nat Med ll(l):56-62. 
59. Zheng X and Bevilacqua PC (2004) Activation of the protein kinase PKR by short doublestranded 
RNAs with single-stranded tails. Rna 10(12):1934-1945. 
60. Sledz CA and Williams BR (2004) RNA interference and double-stranded-RNA-activated 
pathways. Biochem Soc Trans 32(Pt 6):952-956. 
61. Wong JP, Nagata LP, Christopher ME, Salazar AM and Dale RM (2005) Prophylaxis 
of acute respiratory virus infections using nucleic acid-based drugs. Vaccine 23 
(17-18):2266-2268. 
62. Umamaheshwari RB, Ramteke S and Jain NK (2004) Anti-Helicobacter pylori effect of 
mucoadhesive nanoparticles bearing amoxicillin in experimental gerbils model. AAPS 
PharmSciTech 5(2):e32. 
63. Umamaheshwari RB, Jain S, Bhadra D and Jain NK (2003) Floating microspheres bearing 
acetohydroxamic acid for the treatment of helicobacter pylori. / Pharm Pharmacol 
55(12):1607-1613. 
64. Umamaheshwari RB and Jain NK (2004) Receptor-mediated targeting of lipobeads bearing 
acetohydroxamic acid for eradication of helicobacter pylori. / Control Rel 99(l):27-40. 
65. Wong JP, Yang H, Blasetti KL, Schnell G, Conley J and Schofield LN (2003) Liposome 
delivery of ciprofloxacin against intracellular Francisella tularensis infection. / Control 
Rel 92(3):265-273. 
66. Marier JF, Brazier JL, Lavigne J and Ducharme MP (2003) Liposomal tobramycin against 
pulmonary infections of pseudomonas aeruginosa: A pharmacokinetic and efficacy 
study following single and multiple intratracheal administrations in rats. / Antimicrob 
Chemother 52(2):247-252. 
This page is intentionally left blank
INDEX 
131l-lipiodol, 597 
90Y, 408 
99mTc, rhenium-188 (188Re), 595 
99mTc-colloidal nanoparticles, 566 
[3H]thymidine, 486, 487 
3D CT, 569 
5-aminolevulinic acid, 291 
5-fluorouracil, 150, 286 
ABC transporters, 239 
Abelcet, 704 
ACAT, 178 
acetaminohydroxamic acid, 706 
acqueous core, 257, 259, 260, 270 
activated carbon nanoparticles, 589 
active targeting, 500, 502, 503, 507, 516, 51 
adhesion 
monocytes, 15 
adjuvant, 675, 681 
adjuvant treatment, 440 
administration route, 10 
adriamycin, 285 
afferent lymphatic vessel, 552 
agglomeration, 408 
aggregation, 18 
AIDS, 339 
airway geometry, 368, 378, 379 
albumin, 407 
aliphatic polyesters, 30, 31, 33 
allotopic expression, 424 
alum, 676 
AmBisome, 439, 704 
amitriptyline, 536, 541 
amorphous particles, 312, 315 
amoxicillin, 706 
amphiphile, 95-98,100,102-104,110,112, 
113,117 
amphiphilic, 95, 96,100,102 
amphiphilic drugs, 150 
amphiphilic molecules, 125,130 
Amphocil, 704 
amphotericin B, 135,145, 350, 352, 355, 357, 
359, 361, 698 
anatomical, 610 
angiogenesis, 439 
angioplasty, 35 
anionic, 482 
anthracyclines, 437 
anthrax, 556 
anti-HlV drug, indinavir, 580 
anti-infectives, 114 
anti-inflammatory, 361 
anti-prion, 296 
antibacterial, 159, 296 
antibiotics, 359, 360, 377, 385, 388 
antibody, 294, 652, 664, 668 
anticonvulsant, 535 
anticonvulsive, 535 
anticonvulsive activity, 536 
antifungal agent, 352, 355 
antigen release, 35 
antimicrobial, 285 
antimicrobial therapy, 377 
antimony, 698 
antineoplastic, 176,179,181 
antinociceptive effect, 529 
antisense, 700 
antisense oligonucleotides, 337 
AOT, 133,137,150 
apoBlOO, 180 
apoE, 176,180, 535, 541 
apolipoprotein E, 535, 541 
apolipoproteins, 174,176,178,180, 542 
apomorphine, 153 
apoptosis, 340, 429 
aprotinin, 140,141 
aromatic oils, 138 
arsenite, 429 
713 
714 Index 
arterioles, 13,14 
artificial lipoproteins, 174,181 
ascorbyl palmitate, 150 
Aspergillus fumigatus, 701 
asymmetry, 10 
AUC, 147,157 
avidin, 295 
avidin biotin, 577, 586 
avidin-biotin liposome targeting method, 
578 
avidin/biotin-liposome system, 584 
M6F10 tumor cells, 284 
bacterial ghosts, 329 
BBB, 527-530, 533-536, 538-542 
betulinic acid, 429 
bifurcations, 10,13,15,16 
bilayer, 97-100,102,104,110,112,116 
bioavailability, 130,132,135,140-144,146, 
155,157, 282 
biocompatability, 298 
biocompatibility, 403, 509, 517, 619 
bioconjugate, 242 
biodistribution, 33, 37, 530, 538 
biofilm, 159 
biotin, 294, 702 
blastomycosis, 704 
Bleomycin, 588 
blindness, 420 
block copolymers, 30, 31, 33, 36, 96-98,101, 
102,108 
blood velocity, 14 
blood-brain barrier, 34, 239, 312, 527, 534, 
536, 539-542 
bolasomes, 419 
boron neutron capture therapy (BNCT), 
291 
brain, 532-535, 539 
brain concentrations, 531 
brain endothelial cells, 541 
brain perfusion, 532 
brain tumors, 536, 542 
breast cancer, 453 
Brij® 72,532 
bronchial circulation, 380 
Brownian diffusion, 16 
Brownian motion, 14,15 
Brownian relaxation, 405 
buccal delivery, 144 
C-DQAsomes, 428 
caco-2 cells, 297 
calcium phosphate particles, 683 
campthothecin, 542 
camptothecin, 289,531, 532 
cancer, 339,420,437,554 
cancer treatment, 405,411 
Candida albicans, 700 
capillary diameter, 15 
capillary supply, 10 
capsule, 131,140,143 
carbocyanine dyes, 179 
carbon nanotubes, 20, 687 
carbonyl iron, 403 
carboplatin, 283 
cardiomyopathy, 420 
cardiotoxicity, 449 
carmustine, 181 
carrier systems, 697 
catabolism route, 176 
cationic, 482, 494, 495 
cationic lipid, 699 
cavitation, 230 
CD437, 429 
ceftazidime, 703 
cell adhesion, 19 
cell ghosts, 329 
cell monolayers, 13 
cell-penetrating peptides, 2, 5 
cell-specificity, 292 
central nervous system, 527, 532, 535 
ceramide, 429 
characterization, 255, 256, 263 
chemotherapy, 441 
magnetic, 406 
chitosan, 99,100,105,114,116, 627, 652, 
654,657, 658,660, 661, 663,681 
chitosan particles, 683 
chloroquine, 699 
cholesterol, 105,107-112 
cholesterol acyltransferase, 178,180 
cholesteryl esters, 174,177,178,180 
cholesterol-rich emulsions, 174,180 
chylomicrons, 174-176,178,181 
ciprofloxacin, 703 
circadian phase, 530 
cisplatin, 283, 443 
cisplatin-dendrimer complex, 284 
clofazimine, 360 
clonixic acid, 146 
Index 715 
cluster ligands, 293 
CNS, 527, 529,530 
co-precipitation method, 401 
co-solvents, 130,149,150 
coarse-grained model, 420 
coating, 651,652,659,660,662,663,666-668 
cobalt, 402,404 
cochleates, 349-355,357-361 
collective diffusion, 19 
colon-specific delivery, 632 
complex media, 9 
computer tomography (CT), 510 
computed tomography imaging, 561 
confocal fluorescence microscopy, 427 
confocal microscopy, 487 
confusion in reporting lymph node 
delivery, 571 
conjunctiva, 650, 653, 654, 656, 658, 659 
controlled complexation, 618 
convection, 13 
convective flow, 9, 22 
copolymer, 108 
cornea, 650, 653-656, 658, 659, 661, 663, 668 
corneal penetration, 155 
CpG, 686 
CpG-oligonucleotides, 700 
Cremophor EL, 143,146,155,156,158 
Cremophor EL®, 430 
Cryptococcus neoformans, 699 
CT, 569 
CTAB, 688 
curvature, 310 
cyanoacrylic monomer, 259 
cyclosporin A, 130,131,135,140 
cytochrome, 430 
cytoplasm, 9 
cytoskeletal-antigen specific 
immunoliposomes, 486, 492, 495 
cytotoxic agents, 441 
cytotoxicity, 177,178,180,181, 283,425 
dalargin, 530, 541 
Daunoxome, 437 
de Gennes dense packing, 279 
deafness, 420 
decomposition method, 402 
dehydration-rehydration vesicles, 46, 47 
delivery of drugs and vaccines in the small 
intestine, 626 
delivery to the brain, 115 
dendrimer, 11, 278 
dendrimer-antibody conjugate, 294 
dendritic architecture, 278 
dendritic box, 280 
dendritic cell, 30, 37, 330 
dendritic micelles, 279 
dendritic state, 278 
dendroclefts, 282 
dendron, 294 
dendrophanes, 282 
deposition, 11,18 
deposition enhancement factor, 16 
dequalinium salts, 420 
dextran, 701 
dextran-coated, 405 
diabetes, 420 
diagnostic imaging, 382, 383 
diamagnetism, 398 
diaphragm, 586 
diazepam, 157 
dicetylphosphate, 426 
diclofenac, 149,150, 286 
dietary lipids, 175 
diffusion, 10,13, 31, 32, 35 
anisotropic, 20 
collective, 19 
dihydroindolo [2,3-a] indolizine, 699 
dioctanoyl-5-fluoro-2'-deoxyuridine, 532 
discomes, 111, 116 
dissocubes, 314 
dissolution velocity, 308, 309, 317, 322 
DNA, 333 
DNA and protein co-delivery, 51-53 
DNA carriers, 419 
DNA nanoparticles, 688 
docetaxel, 293 
domain structure, 400 
magnetic, 400 
dosage form, 132,134,140,146,154, 634 
double microemulsions, 133-135 
double-targeting, 429 
Doxil, 437 
doxorubicin, 113,114, 289, 407, 447, 530, 
531,533, 534,536-539,541,542 
doxorubicin (Dox), 177,178 
DQAplexes, 425 
DQAsomes, 419 
droplet, 126,127,129-133,135,144-147, 
150,151,155,156,159 
drug, 255-258, 261-263, 265-271 
716 Index 
drug absorption, 127,130,142,154 
drug carrier, 349, 350, 361 
drug carrier nanoparticle, 9 
drug delivery, 58-62, 64, 68, 70, 72, 76, 79, 
80, 95, 98,104,106,110,113, 282, 329, 349, 
352, 463,464, 472-474, 476, 481, 482, 485, 
492,495 
magnetic, 406 
drug delivery system, 127,130-132,134, 
144,147,160, 437 
drug delivery to lymphatics, 553 
drug efflux pumps, 34 
drug release, 30, 32, 37 
drug-loaded nanoparticles, 615 
drug-loaded tumor cell system (DLTC), 340 
dry emulsions, 134,135 
dry powder inhaler, 369-371, 373, 382 
duodenal administration, 531 
E-mediated lysis, 330 
early endosomal release, 425 
EEC, 535 
efferent lymphatic vessels, 552 
elasticity 
vesicles, 12 
electrostatic repulsion, 19 
embolization, 405, 408, 505 
emend, 317 
emulsification-solvent-evaporation, 617 
emulsifying wax/Brij® 78, 532 
emulsion, 125-129,131,132,134,135,145, 
147,157, 322-324, 697 
emulsion polymerization, 31, 32 
emulsion-solvent-diffusion, 617 
encapsulation, 282 
encephalitis, 534 
endocytotic uptake, 540-542 
endolymphatic radioisotope therapy, 591 
endothelial, 501, 503,504,506-511,513, 
516-518 
endothelial cell, 527, 538-541 
enhanced capillary permeability, 697 
enhanced permeability and retention, 283 
enhanced permeability and retention effect 
see EPR effect, 2,4, 6,409,439,442,446, 
11, 502, 503,507,512 
enkephalin, 529 
enterocytes, 10 
enzyme, 500, 508, 509, 511-516 
enzyme delivery, 515, 516 
enzyme replacement therapy, 511 
enzyme therapy, 511, 514 
Epicuron® 200,531 
epifluorescence, 487 
EPR, 283 
erosion, 33,35 
erythrocyte ghosts, 329, 333 
erythrocytes, 13 
Escherichia coli, 700 
etoposide, 286 
excipients, 131,132,136,137,148,156 
experimental acute myocardial infarction, 
482 
extravasation, 9, 446 
extrusion, 12 
eye, 649-652, 655, 657-659, 662, 665-668 
fast action onset, 318 
fate of nanoparticles in lymph nodes, 561 
ferritin, 397 
ferrofluid, 401 
ferromagnetism, 398 
ferrosilicone, 408 
fetuin, 482 
filaria, 555 
flow 
GI tract, 11 
fluorescence microscopy, 331 
fluorescent quantum dots, 568 
fluorocarbon, 126,158,159 
folate, 148 
folic acid, 292 
folic receptor, 292 
follicle-associated epithelium (FAE), 613 
Francisella tularensis, 706 
freeze-fracture, 354 
functionalized, 651, 652, 663, 668 
Fungizone, 698 
gamma scintigraphic imaging, 482 
gastric irritancies, 318 
gastric mucosa, 625 
gastro-intestinal tract, 37, 609 
gelatin, 131,133 
gene delivery, 105,175,179, 506, 507, 510 
gene delivery carriers, 179 
gene therapy, 338, 367, 386-388 
gene transfer, 409 
genetic vaccination, 43 
gentamicin, 702 
Index 717 
GI fluid, 131 
glioblastoma, 536-538 
glioblastoma 101/8, 542 
glycodendrimers, 293 
glycol chitosan, 99 
gold-198 colloid, 597 
Gram-negative bacteria, 330 
haematocrit, 13 
haemolytic effect, 298 
hand-foot syndrome, 452 
heart concentration, 531, 534 
heart toxicity, 531 
Helicobacter pylori, 705 
hemolysis, 357 
heparin, 141,142, 704 
hexadecyl diglycerol ether, 111, 112 
hexapeptide dalargin, 529, 541 
hFR (high affinity folate receptor), 292 
high pressure homogenization, 188,189, 
191,192 
high-density lipoprotein (HDL), 174,175, 
178-181 
histone, 426 
HIV, 295,555,580 
homogenization, 314-316, 322, 323 
human colon cancer, 431 
humidity, 368, 371, 373, 378, 379 
hyaluronic, 652, 657, 659, 660 
hydrodynamic bridging, 22 
hydrodynamic diameter, 100,101 
hydrogels, 30, 35 
hydrophilic, 95, 97, 98,102-106,108,110, 
112,115,116 
hydrophilic spacer, 98 
hydrophobic, 95-106,108 
hydrophobic drags, 142,151,153,173,175 
hydrophobic interactions, 19 
hydrosols, 313 
hyperthermia 
magnetic, 405 
hypotonic dialysis, 334 
1-131 (131I)-lipiodol, 591 
ibuprofen, 287 
ICso, 297 
idarubicin, 531, 542 
IgE, 295 
imaging, 510, 697 
imaging agents, 113,115,117 
immiscible phases, 126 
immune modulation, 597 
immune response, 557 
immunization, 154 
immunoliposomes, 481^488, 492, 495 
immunostimulatory oligonucleotide, 686 
In-Ill, 482 
increases in survival times (1ST), 537 
IND, 295 
indomethacin, 155, 287 
infectious diseases, 697 
infertility, 420 
influenza A virus, 705 
influenza vaccine, 51, 53 
inhalation toxicity, 404 
inspiratory flow rate, 371-373, 375 
insulin, 140,141 
integrin, 242 
interactions 
with blood components, 16 
interfacial polymerization, 256, 257, 
259-261, 265,266,268,269 
intermediate-density lipoprotein (IDL), 
174 
internalization, 177,178,180 
intralymphatic drug delivery, 587 
intramuscular delivery, 144 
intranasal administration, 384, 387, 388 
intraocular, 649, 650, 653-655, 657, 659, 660, 
665-668 
intraoperative radiotherapy, 
593 
intraperitoneal, 581 
intraperitoneal clearance, 583 
intraperitoneal liposome encapsulated 
drugs, 582 
intrapleural injection, 585 
intratumoral administration, 595 
intratumoral injection, 22, 23 
intratumoral radionuclide therapy, 593 
intravaginal drug delivery, 156 
intravenous delivery, 144,147 
intravitreal, 649-651, 665-667 
iodinated nanoparticles, 561, 562 
iron, 402 
iron oxide particle, 701 
iron-carbon, 407 
isoniazid, 703 
isotropic, 129,131,132,157 
IV injection, 13 
718 Index 
jet-stream, 314 
Kaposi's sarcoma, 289, 453 
KB cells, 293 
Kelvin equation, 310-312 
Klebsiella pneumoniae, 702 
Kupffer cell, 13, 33 
kyotorphin, 534, 541 
Labrasol, 141,142,149,150,157 
Langendorff, 488, 489 
LDL-receptor, 173,177-181 
lecithin, 133-135,138,140,145,146,151, 
159, 427 
lectin, 664, 668 
leishmaniasis, 698 
leukocytes, 21 
lidocaine, 132,144,151-154 
lipid based formulation, 180 
lipid-coated, 225 
lipofectin, 425 
lipolysis, 176 
lipophilic drug release, 223 
lipoprotein receptors, 541, 542 
lipoproteins, 173-181 
liposome, 12, 43, 45-52, 353-355, 360, 361, 
437,481^84, 486-488, 492-495, 574, 697 
liquid crystalline phase, 128 
liquid crystalline state, 110 
liver, 10 
liver cancer, 407 
liver cell targeting, 176 
loading capacity, 620 
long-circulating liposomes, 442, 701 
long-circulating nanoparticles, 34 
long circulating microemulsions, 147 
long-term survivors, 539 
long-time survivers, 537 
lonidamine, 429 
loperamide, 534, 541 
low density lipoprotein (LDL), 173-178, 
181 
lung cancer, 382 
lung clearance, 378, 379 
lungs, 499-504, 507, 508, 510, 511 
lymph node anti-infectious agent delivery, 
580 
lymph node retention efficiency, 573 
lymph node targeting method, 577 
lymph nodes, 10, 35,405,551,587 
lymph vessels, 10,11 
mesenteric, 11,12 
lymphatic circulation, 379 
lymphatic clearance, 560 
lymphatic radiotherapy, 592 
lymphatic system, 175, 379, 380, 382-384, 
549 
lymphatic vessels, 550 
lymphatics, 10, 35 
lymphoscintigraphy, 382, 383 
lyophilization, 313 
M-cells, 10,11,36, 37 
macromolecules, 377, 378, 380 
macrophage phagocytosis, 561 
macrophages, 5, 30, 33-35, 37, 331, 336, 368, 
380,381,383,385,387, 534 
maghemite (Fe2C>3), 401 
magnetic drug delivery, 406 
magnetic field, 398,410 
magnetic microspheres 
albumin, 407 
general, 397 
iron-carbon, 407 
magnetic nanoparticles, 563 
biocompatibility, 403 
encapsulation, 403 
general, 397 
toxicity, 403 
magnetic properties, 398 
magnetic response 
temperature dependence, 400 
magnetic resonance (MRI) contrast agents, 
563 
magnetic resonance imaging (MRI), 410 
magnetic targeting device, 410 
magnetite (Fe304), 399,401 
magnetization curve, 399 
magnetofection, 409 
magnetotactic bacteria, 397 
Mannose, 699 
mass transport, 19 
massage, 560, 578 
Massart's method, 401 
material properties, 400 
mechanism, 539, 540 
mediastinal lymph node, 585 
mediastinal lymph node targeting, 586 
mefenamic acid, 287 
melting point depression method, 132 
Index 719 
membrane, 102,106-112 
metastable, 96 
metastases, 439 
metastatic lymph nodes, 589 
methotrexate, 285, 590 
methylprednisolone, 289 
micelles, 32, 57-80, 97,102,104,108, 
110-112,128-130,133,144, 574, 697 
microbubbles, 225 
microemulsion, 125-160, 531 
microemulsion gels, 133,134 
microfluidizer, 314, 316 
microorganism, 697 
microparticles, 678 
microstructure, 128-130,139,151,152 
microtubules, 15 
migraine, 420 
migration 
gravity induced, 19 
shear-induced, 14 
viscosity induced, 15 
mitochondria, 5, 419 
mitochondrial gene therapy, 424 
mitochondrial membrane potential, 425 
mitochondrial protein import machinery, 
424 
mitochondrial size, 490, 491 
mitomycin C, 590 
molecular imprinting, 282 
molecular recognition, 293 
molecular weight, 96-98,100-103,108,117 
mononuclear phagocyte system, 697 
Monte Carlo simulations, 420 
MRI imaging, 563 
mtDNA, 424 
mucoadhesive, 657, 659-664 
mucociliary clearance, 368, 379, 380 
mucosa, 142,144,158, 655-659, 661, 662, 
664, 668 
mucosal, 682 
mucosal adjuvant, 684 
mucosal vaccines, 682 
mucus, 657, 659, 662-664 
multi-drug resistance, 429 
multi-prodrug, 289 
multilamellar, 482 
multivalency, 293 
Mycobacterium tuberculosis, 385-387, 703 
Myocet, 437 
Neel relaxation, 405 
naloxone, 529 
nano-lipid vesicles, 481, 482 
nano-scaffolding, 279 
nanocapsule, 213-215, 218-220, 222, 
255-271 
nanocarrier, 499, 500, 502-509, 511-513, 
515-518, 697 
nanocochleates, 352, 354, 355, 360 
nanocrystals, 20, 307-313, 315, 317-324 
nanoEdge, 315, 319 
nanoemulsion, 126,127,129,131,174,176, 
177 
nanoerythrosomes, 337 
nanomedicines, 439 
nanoMorph, 313 
nanonization, 308, 317 
nanoparticle 349, 350, 352, 354, 368, 375, 
381-388,481, 483^85,495, 528,530,532, 
534,535,538, 540,541, 549, 580, 697 
active targeting, 471, 472, 476 
receptor-mediated endocytosis, 
469, 471 
application in, 464, 473 
acquired immune deficiency 
syndrome (AIDS), 464 
gene therapy, 476 
leishmaniasis, 464, 468, 471, 473, 
474,476 
pulmonary tuberculosis, 464, 473 
trypanosomiasis, 464, 474 
chemotherapy, 473, 477 
definition, 400 
surface modifications, 469 
uptake, 464 
factors influencing, 468 
mechanism of, 465 
sites of, 464 
nanoparticle diagnostic agents, 565 
nanoparticle flow, 9 
nanoparticle lymph node drug delivery, 
571 
nanoparticle size, 559 
nanoparticle surface, 559 
nanoprecipitation, 31, 32 
nanoPure, 315, 316, 318 
nanoscale container, 279 
nanosphere, 30-32, 34-37 
nanosuspension, 307-309, 314-324 
nanotechnology, 481, 495 
720 Index 
naproxen, 288 
nasal mucosa, 37 
nasal route, 157 
nebulizer, 369-371, 382, 383, 386 
Neobee M-5,143 
neurodegenerative diseases, 420 
neuromuscular diseases, 420 
neutral, 482 
nickel, 404 
nifedipine, 288 
Niobe system, 410 
niosomes, 12, 95-98,104,108,110-117 
nitrendipine, 142 
nitroblue tetrazolium, 489^92 
nociceptive reactions, 530 
non-pyrogenic, 619 
non-reactive, 619 
non-specific endocytosis, 425 
non-steroidal anti-inflammatory drug 
(NSAID), 287 
nonionic, 133,136,137,142,153,154 
nonionic surfactant, 133,136,137,142,153 
Noyes-Whitney, 309 
nubiotics, 698, 705 
nystatin, 700 
O-stearylamylopectin, 703 
ocular, 649-664, 666-668 
ocular administration, 37 
ocular delivery, 297 
oil-to-water ratios, 128 
oily core, 256,262,265, 270 
oleic acid, 141,147,151,152 
ophthalmic dosage forms, 154 
opsonization, 33, 532 
optical, 568 
optical imaging, 179 
oral administration, 536, 610 
oral cavity, 621 
oral delivery, 350, 353-355, 360, 361 
oral drug delivery, 139 
osmotic lysis, 333 
ovarian cancer, 453 
oxaliplatin, 283 
oxidation 
controlled, 402 
P-glycoprotein (Pgp), 528, 532, 534, 539 
P-glycoprotein efflux pump, 297 
P-glycoprotein pumps, 613 
P-gp efflux, 297 
PACA, 652-655,659,660, 662, 667 
paclitaxel, 146,147,181, 289, 430, 437, 532, 
542 
palmar-plantar erythrodysesthesia, 452 
PAMAM, 278 
paramagnetism, 398 
parenteral delivery, 296 
particle bridging, 22 
particle diameter, 10 
particle flow, 10 
particle migration, 14 
particle shape, 20 
particle size, 309-313, 315, 316, 369, 
372-376, 378, 379, 383, 385 
particle stability, 24 
particle-image velocimetry, 15 
passive targeting, 502, 503, 505 
PathFinder, 312 
patient compliance, 318 
pDNA-MLS peptide conjugates, 427 
PECL, 653-657, 660-663 
PEG, 532,533, 660, 662, 663, 667 
PEG 2000, 533 
PEG-liposomes, 702 
PEG-PHDCA, 533,534 
PEG-spacer, 532 
PEGylated [14C]-poly[methoxy poly 
(ethylene glycol) 
cyanoacrylate-co-hexadecyl 
cyanoacrylate], 533 
pegylated liposomal doxorubicin, 448 
pegylated liposomes, 442 
PEGylated solid lipid nanoparticles, 534 
PEGylation, 298, 532, 542 
pegylation, 443 
penicillin, 295 
penicilloylated dendrimers, 295 
peptides, 140,141,149,159 
permeability, 620 
Peyer's patches, 10,11, 36, 612 
pGL2, 493 
phagocytic cells, 331 
pharmacokinetics, 357-359, 698 
Pharmasol, 315, 316, 318 
phase diagram, 128,136 
phase transition temperature, 112 
phosphatidylserine, 425, 699 
phospholipid, 95,116, 350-352 
phospholipid nanoemulsion, 176,177 
Index 721 
photodynamic therapy (PDT), 291 
physiological, 610 
piperine, 699 
piston-gap, 314-316 
PLA, 654, 660, 662, 666 
plasmid, 330 
platelet flow, 21 
PLGA, 146,147, 653, 655, 665-667 
PLGA nanoparticles, 573 
PLGA, poly lactic acid (PLA), and poly 
(fumaric anhydride-co-sebacic 
anhydride) have, 628 
Pluronic® F68, 531 
pluronics, 137 
poloxamer, 33, 534 
poloxamer 188, 531, 704 
poloxamer 908, 533 
poloxamine, 33 
poloxamine 908, 532, 534, 540 
Poly (DL-lactide-co-glycolide) (PLGA), 
poly (e-caprolactone) (PCL), poly 
(alkylcyanoacrylates), poly 
(styrene-co-maleic anhydride), poly 
(divinylether-co-maleic anhydride), poly 
(vinyl alcohol), poly (ethylene glycol), 
615 
poly (hexadecyl cyanoacrylate) 
nanoparticles, 533 
poly (isobutylcyanoacrylate), 627 
poly (lactide) (PLA), 678 
poly (lactide-co-glycolide) (PLG), 678 
poly(alkylcyanoacrylate), 31, 34 
poly(alquilcyanoacrylate), 654, 660 
poly(amidoamine), 278 
poly(butyl cyanoacrylate), 529, 530, 
534-536,538,541 
poly(epsilon-caprolactone) nanospheres, 
704 
polyethylene glycol) (PEG), 2, 5,105,110, 
504, 516 
polyethylene oxide), 102,104,110 
poly(ethyleneglycol), 31, 33, 660, 662 
poly(ethylenimine), 100,105 
poly(ethylenimine) amphiphiles, 100 
poly(ethylimide) (PEI), 511, 514 
poly(glutamic acid), 426 
poly(L-lysine), 98,100,105 
polyflactic acid), 654, 660 
poly(lactic-co-glycolic acid), 514 
poly(lactide-co-glycolide), 30 
poly(lysine), 295 
poly(methyl methacrylate), 530 
poly(n(2-hydroxypropyl)methacrylamide) 
(HPMA), 514 
poly-epsilon-caprolactone, 652, 654, 655, 
660 
polyacrylamide, 98, 99 
poly disperse nanoparticles, 13,15 
polyethyleneglycol, 482 
polyhedral vesicles, 20,111-113 
polylCLC, 705 
polylactic acid, 534 
polymer, 57-74, 76-80, 255-266, 269-271, 
381, 385-387 
polymeric nanoparticles, 29, 30, 32-34, 36, 
37, 609 
polymeric vesicles, 95-98,102-108,117 
polymerization, 97, 98,102-104,106 
polymersomes, 101,102,104 
polymorphonuclear leucocytes, 340 
polyplexes, 510,514 
polysorbate 20,535 
polysorbate 20, 40, and 60, 534 
polysorbate 80, 529-541 
polysorbate 80-coated nanoparticles, 536 
polysorbate 80-coated poly(butyl 
cyanoacrylate) nanoparticles, 537, 539, 
542 
polystyrene latex, 12 
polystyrene nanoparticles, 35, 36 
poorly soluble, 307, 308, 311, 313, 321, 322, 
324 
poorly soluble pharmaceuticals, 3 

pores, 23 
particle entry, 23 
positive tumor margins, 593 
Positron Emission Computed Tomography 
(PET), 570 
precipitation, 309, 313, 315 
pressurized metered dose inhaler, 369-372, 
376, 381 
pro-apoptotic drugs, 429 
probenecid, 535 
prodrug, 289 
protein E, 330 
proteins, 368, 377, 378, 380 
Pseudomonas aeruginosa, 705 
pSV-^-gal vector, 493^95 
pulmonary, 499-503, 505-513, 516-518 
pulmonary delivery, 158 
722 Index 
pulmonary drug delivery, 501, 509 
quantum dots, 11,177 
quinolinium derivatives, 419 
radioisotopes 
diagnostic, 408 
therapeutic, 408 
rapamune, 308, 317 
RAST (radioallergosorbent test), 295 
receptor-mediated, 542 
rectal drug delivery, 156 
removal of particles, 18 
resistant, 703 
respiratory syncytial virus, 387, 388, 705 
Responsive Release - pH, 106 
Responsive Release - Temperature, 111 
Reticulo-endothelial system (RES), 404, 
443,445,469, 532, 533, 
bone marrow, 464-470 
disorders, 463, 464, 473, 476 
infectious, 464 
non-infectious, 464 
liver, 465-470, 474^76 
lymph node, 469, 470 
macrophage, 464^72, 474, 476 
monocytes, 464, 466, 467, 469, 474 
spleen, 464-169, 473^76 
retina, 653, 665, 667 
reverse biomembrane vesicles, 338 
Reynolds number, 16 
rhenium-186, 595 
rhenium-labeled liposomes, 595 
rheology, 13,14 
rifampicin, 703 
RNA, 333 
RNA interference, 700 
route of administration, 10 
rugosity 
of surfaces, 18 
salting out, 617 
saturation solubility, 309-313, 322 
segregation, 10 
self assembly, 95-98,100-103,108 
self-association, 423 
self-diffusion, 15 
self-emulsification, 131 
self-emulsifying drug delivery system, 131 
self-immolative dendrimers, 289 
sensitivity to, 620 
sentinel lymph node, 565 
sentinel lymph node identification, 566 
shear 
radial variation, 13 
shear forces, 9,10 
shear-induced migration, 14 
sialoglycoprotein, 482 
silver salts, 285 
single photon emission computed 
tomography (SPECT)/computed X-ray 
tomography (CT) systems, 569 
siRNA, 698, 700 
skin, 134,148-154 
SMBV nanoparticles, 683 
soft contact lens, 156 
SolEmuls, 322-324 
solid lipid nanoparticles (SLN), 187-205, 
531-533,542 
solid liquid nanoparticles, 152 
solid-state emulsions, 135 
solubility enhancement, 288 
Solulan C24,109-113 
solvent displacement, 31 
solvent displacement method, 680 
solvent evaporation, 31 
somatic mutations, 420 
sonication, 136 
Soy phosphatidylserine, 352 
SPECT/CT, 570 
Speiser, 528 
spermidine, 426 
spheroids 
oblate, 20 
SPI-77, 456 
spinal cord, 534 
SPION, 404 
spleen, 10,33, 34 
splenotropic, 34 
stability, 98,102,104,110 
state diagrams, 19 
STD, 295 
stealth, 442, 499, 504, 507, 512, 513, 516 
stealth liposomes, 446 
stealth nanoparticles, 34 
stealth vectors, 214 
stem cells, 338 
sterically-stabilized, 442 
streptokinase, 409 
structure activity relationship studies, 422 
Index 723 
subcellular localization, 503 
subconjunctival, 649, 665 
subcutaneous injection, 13 
sulphur colloid, 13 
supercooled melts, 190,198,199, 201 
supercritical fluid technology, 618 
superparamagnetism, 400 
surface ligands, 18 
surface properties, 534 
surface receptors 
interaction, 18 
surfactant, 125-129,131,133-139,141-146, 
149-159, 309, 535 
sustained release, 34, 35, 329 
swellability, 620 
swollen micelles, 129,130 
SYBR™ Green I, 425 
systemic adjuvant, 686 
systemic circulation, 380, 381 
systemic delivery, 296 
tail-flick test, 529 
targeted delivery, 148, 294 
targeting, 104,105,110,113,115,117, 663, 
664, 668 
tat peptide, 409 
tat-CLIO, 409 
technegas, 382 
terminal filtration, 130 
test, 529 
tetanus toxoid, 384, 385 
therapeutic applications, 255 
therapeutic proteins, 140,141 
thermoresponsive, 111, 112 
thiamine, 532 
thrombolysis, 409 
tight junction, 528, 539, 540 
tissue distribution, 698 
tobramycin, 360, 531, 542, 706 
tocopheryl polyethylene glycol 1000 
succinate (TPGS), 141 
topical, 649-652, 655-658, 660, 665, 668 
topical delivery, 148 
toroids, 20 
toxicity, 38, 402, 449, 538 
toxicology and regulatory aspects, 636 
transcytosis, 542 
transdermal delivery, 139,148,150,152, 297 
transport 
nanoparticles, 10 
triglycerides, 135,137,139,141-144,150, 
151,174,176 
triton-X 100,107 
trypan blue, 486, 487 
tuberculosis, 556 
tubes, 20 
tubocurarine, 535, 541 
tubules 
multi-bilayer, 12 
tuftsin, 700 
tumor, 34 
tumor cells, 340 
tumor diagnosis, 179 
tumor interstitium, 10 
tumor necrosis factor-alpha, 700 
tumor therapy, 595 
tumor vessel diameter, 22 
Tween, 134,137,138,141-143,145,146,149, 
151,153,154,157,159 
Tween® 80, 529 
ultrasound nanobubbles, 569 
unilamellar, 482 
unimolecular encapsulation, 279 
upper GI malignacies, 589 
urokinase, 409 
UV irradiation, 150 
vaccine, 333, 383-385, 387,388, 675,682 
valproic acid, 536 
van der Waals forces, 19 
vancomycin, 703 
vascular thrombosis, 237 
vasculature, 499-503, 505-508, 511, 513, 
517,518 
vector molecules, 4 
venlafaxine, 287 
ventilation scan, 382 
venules, 13,14 
very low density lipoprotein (VLDL), 
174-178,181 
vesicle, 95-108,110-113,115-117, 350, 352, 
354, 361 
vesicle formation index, 100 
vesicle size, 102,113 
vincristine, 147 
vinorelbine, 429 
viral-sized colloids, 13 
viscoelasticity, 620 
viscosity, 127,156,157 
724 Index 
vitamin E, 132,147 water-soluble drug, 134,143,149,152,154 
vitreous, 657, 665-667 wheat germ agglutinin, 632 
VivaGel™, 295 
VP-16, 429 zeta potential, 147,151