Bioinspired and Biomimetic Polymer Systems for Drug and Gene Delivery 9783527334209, 9783527672738, 9783527672721, 9783527672714, 9783527672752, 3527334203

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Edited by Zhongwei Gu Bioinspired and Biomimetic Polymer Systems for Drug and Gene Delivery

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Edited by Zhongwei Gu

Bioinspired and Biomimetic Polymer Systems for Drug and Gene Delivery

Editor Prof. Zhongwei Gu

National Engineering Research Center for Biomaterials Sichuan Univ. Wangjiang Road 29 610064 Chengdu China

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V

Foreword In recent years, the rapid development of polymer science and advances in modern medicine, pharmacy, biology, and engineering have fostered the emergence of a new field focused on the theory and technology underlying drug delivery. This inter-disciplinary field is called drug delivery systems (DDS). It shows great promise and has become a hotspot in biomedical material research, especially in biomedical polymers. The successful development of advanced, efficient DDS depends on the design and construction of the materials and micro devices involved. The research frontier focuses mainly on targeted delivery, especially cell and molecular targeting, and on controlled release stimulated by the tissue or cellular microenvironment. The complex in vivo physiological and pathological environment often obscures the effects of active targeting. In this way, producing a highly efficient system capable of active targeting in vivo is the key to improving the efficacy of DDS. Drug release systems capable of biological sensing are called bioinspired and biomimetic delivery systems. They automatically adjust the drug release in response to external stimuli, such as changes in temperature, pH, magnetic fields, ultrasound, and electric fields. They have received a considerable amount of attention from researchers and pharmaceutical companies worldwide. Drug release systems that can be switched on and off via self-feedback upon changes in the chemical or physical signals given off by a lesion or intelligent carrier have drawn particular interest. Systems that can undergo rapid stimuli-responsive controlled release under in vivo microenvironment conditions would be far more useful to actual clinical treatment regimens. This book embodies the wisdom and achievements of renowned experts and research teams in this field from China, the United States, Germany, Japan, and Korea. The discussion provided herein covers the most important, active, and cutting-edge parts of this field, reflecting the latest developments and trends in DDS research. The chief editor, Professor Zhongwei Gu, studied under the pioneer biomedical polymers in China–Professor Xin-De Feng (Academician of Chinese Academy of Sciences). Gu entered this field in the 1970s and has become a wellknown professor of polymer biomaterials in China. It is our hope that this book will promote scientific research and biomedical applications in the vibrant and

VI

Foreword

exciting area. Young academics and professionals interested in DDS may also benefit from this treatise. We would like to thank all our editors for their hard work and dedication. We would also like to thank John Wiley & Sons Publishing Company and Chemical Industry Press for their forward-looking strategic vision and the timely publication of this book. August 2014

Professor Ren-Xi Zhuo Academician of Chinese Academy of Sciences IUS-BSE Fellow Wuhan University Wuhan, China

VII

Contents List of Contributors XIII Preface XIX 1

Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics 1 Jiyuan Yang and Jindˇrich Kopeˇcek

1.1 1.2 1.2.1 1.2.2 1.2.2.1

Introduction 1 Water-Soluble Polymers as Carriers of Anticancer Drugs 2 First Generation Conjugates – Design, Synthesis, and Activity 2 Analysis of Design Factors That Need Attention 2 Design of Conjugates for the Treatment of Noncancerous Diseases 2 Combination Therapy Using Polymer-Bound Therapeutics 3 New Targeting Strategies 4 Relationship between Detailed Structure of the Conjugates and Their Properties 5 Impact of Binding a Drug to a Polymer on the Mechanism of Action 6 Mechanism of Internalization and Subcellular Trafficking 7 Relationship between the Molecular Weight of the Carrier and the Efficacy of the Conjugate 7 Design of Second Generation Conjugates – Long-Circulating and Backbone Degradable 8 RAFT Copolymerization for the Synthesis of Conjugates 8 Click Reactions for Chain Extension into Multiblock Copolymers 10 Biological Properties of Long-Circulating Macromolecular Therapeutics 10 Summary of Part 2 and Future Prospects 14 Drug-Free Macromolecular Therapeutics – A New Paradigm in Drug Delivery 15 Biorecognition in Hybrid Polymer Systems 15 Coiled-Coils in Biomedical Systems 16

1.2.2.2 1.2.2.3 1.2.2.4 1.2.2.5 1.2.2.6 1.2.2.7 1.2.3 1.2.3.1 1.2.3.2 1.2.3.3 1.2.4 1.3 1.3.1 1.3.2

VIII

Contents

1.3.3 1.3.4 1.4

Coiled-Coil Based Drug-Free Macromolecular Therapeutics: Design, In Vitro, and In Vivo Activity 17 Potential, Limitations, and Future Prospect of Drug-Free Macromolecular Therapeutics 18 General Summary and Outlook 20 Acknowledgments 21 References 21

2

Dendritic Polymers as Targeting Nanoscale Drug Delivery Systems for Cancer Therapy 29 Kui Luo and Zhongwei Gu

2.1 2.2

Introduction 29 Functional Dendritic Polymers Based Drug Delivery Vehicles for Targeting Tumor Therapy via EPR Effect 30 Functional Dendritic Polymers for Encapsulation of Anticancer Drugs 32 Chemical Conjugation Functional Dendritic Polymers as Drug Delivery Systems 37 Tumor Targeting Moieties Functionalized Dendritic Drug Delivery Vehicles for Cancer Therapy 45 Conclusion 54 References 54

2.2.1 2.2.2 2.3 2.4

3

Composite Colloidal Nanosystems for Targeted Delivery and Sensing 61 Pilar Rivera Gil, Moritz Nazarenus, and Wolfgang J. Parak

3.1 3.1.1 3.1.2 3.2 3.3 3.3.1 3.3.2 3.4 3.4.1 3.4.1.1 3.4.1.2 3.4.2 3.5

Introduction 61 Working Toolkit 62 Engineering a Multifunctional Carrier 63 Objective 66 Cellular Behavior of the Carrier 66 Intracellular Fate 66 Biocompatibility 69 Applications 71 Delivery with Multifunctional PEM Capsules 71 Magnetic Targeting and Magnetofection 71 Strategies for Controlled Opening 73 Intracellular Ion Sensing 75 Conclusions 77 Abbreviations 77 References 78

4

Polymeric Micelles for Cancer-Targeted Drug Delivery Huabing Chen, Zhishen Ge, and Kazunori Kataoka

4.1

Introduction 85

85

Contents

4.2 4.3 4.4 4.5 4.6 4.7 4.8

Micelle Formulations in Clinical Development 85 Particle Size of Micelles 89 Morphology of Micelles 92 Targeting Design of Micelles for Enhanced Accumulation and Cell Internalization 94 Functional Designs of Micelles 96 Design of Micelles for Gene Delivery 99 Challenge and Future Perspective 103 References 104 109

5

Biomimetic Polymers for In Vivo Drug Delivery Wenping Wang and Kinam Park

5.1 5.2

Introduction 109 Commonly Used Biomimetic Polymers and Their Applications in DDS 110 Polylactones and Their Modifications 110 Poly(lactic acid) (PLA) 110 Poly(lactic-co-glycolic acid) (PLGA) 113 Poly(ε-caprolactone) (PCL) 118 Dendrimer 124 Structure and Properties of Dendrimers 124 Types of Dendrimers 124 Applications of Dendrimers as Carriers in Drug Delivery Systems 124 Synthetic Polypeptides 134 Challenges and Perspectives 135 References 136

5.2.1 5.2.1.1 5.2.1.2 5.2.1.3 5.2.2 5.2.2.1 5.2.2.2 5.2.2.3 5.2.3 5.3

6

Drug Delivery from Protein-Based Nanoparticles 149 Dan Ding and Xiqun Jiang

6.1 6.2 6.2.1 6.2.2 6.2.3 6.2.4 6.3 6.3.1 6.3.2

Introduction 149 Preparation of Protein-Based Nanoparticles 150 Desolvation 150 Emulsification 151 Coacervation 151 Polymer–Monomer Pair Reaction System 151 Drug Delivery from Albumin-Based Nanoparticles 152 Albumin-Based Nanoparticles as Drug Carriers 152 Targeting Ligand-Functionalized Albumin-Based Nanoparticles 154 Nanoparticle Albumin-Bound (nab) Technology 156 Drug Delivery from Gelatin-Based Nanoparticles 156 Gelatin-Based Nanoparticles as Drug Carriers 158 Targeting Ligand-Functionalized Gelatin-Based Nanoparticles 160 Site-Specific Drug Delivery System 162

6.3.3 6.4 6.4.1 6.4.2 6.4.3

IX

X

Contents

6.5

Drug Delivery from Other Protein-Based Nanoparticles 163 References 165

7

Polymeric Gene Carriers 171 Xuesi Chen, Huayu Tian, and Xiuwen Guan

7.1 7.1.1 7.1.1.1 7.1.1.2 7.1.1.3 7.1.1.4 7.1.2 7.1.2.1 7.1.2.2 7.1.2.3 7.1.2.4 7.2 7.2.1 7.2.1.1 7.2.1.2

Gene Therapy and Gene Carriers 171 Gene Therapy 171 The Concept of Gene Therapy 171 Development and the Present Situation of Gene Therapy 171 Methods and Strategies of Gene Therapy 172 Research Contents and Challenges of Gene Therapy 174 Gene Carriers 175 The Concept of Gene Carrier 175 The Necessity of the Gene Carrier 175 Requirements of Gene Carrier 176 Classification of Gene Carrier 176 Polymeric Gene Carriers 178 Cationic Polymer Gene Carriers 178 Process of the Polycation Vector Mediated Gene Delivery 179 Categories and Research Situation of the Cationic Polymer Gene Vector 180 PEI Grafting Modification Polymeric Gene Carriers 183 Amino Acid Derivatives Modified Polymeric Gene Carriers 183 Poly(glutamic acid) Derivatives Modified PEI 184 Polyphenylalanine Derivatives Modified PEI 186 PEG Modified Hyperbranched PEI 187 Low Molecular Weight (LWM) PEI Base Polymeric Gene Carriers 188 Crosslinked Polycations 188 Crosslinked Polycation OEI-CBA 188 Crosslinked Polycation OEI-PBLG-PEGDA 189 Hexachlorotriphosphazene Crosslinked Polycation 190 Grafted Polycations 190 Grafted Cationic Polymer MP-g-OEI 190 Graft Cationic Polymer N-PAE-g-OEI 191 Graft Cationic Polymer mPEG-b-PMCC-g-OEI 192 Targeted Shielding System for Polymeric Gene Carriers 192 Static Shielding System 192 Poly(glutamine acid) Shielding System and PEGylations 195 Sulfonamides Related Shielding System 195 Other Design Strategies of Cationic Gene Carrier 196 Conclusion 197 References 197

7.3 7.3.1 7.3.1.1 7.3.1.2 7.3.2 7.4 7.4.1 7.4.1.1 7.4.1.2 7.4.1.3 7.4.2 7.4.2.1 7.4.2.2 7.4.2.3 7.5 7.5.1 7.5.1.1 7.5.1.2 7.5.2 7.6

Contents

8

pH-Sensitive Polymeric Nanoparticles as Carriers for Cancer Therapy and Imaging 203 Yi Li, Guang Hui Gao, Ick Chan Kwon, and Doo Sung Lee

8.1 8.2 8.2.1 8.2.2 8.2.3 8.3 8.3.1 8.3.2 8.3.3 8.3.4 8.3.5 8.4 8.5

Introduction 203 pH-Sensitive Polymers 204 pH-Sensitive Anionic Polymers 205 pH-Sensitive Cationic Polymers 207 pH-Sensitive Neutral Polymers 208 pH-Sensitive Polymers as Drug Carriers 209 pH-Sensitive Polymer–Drug Conjugates 210 pH-Sensitive Polymeric Micelles 210 pH-Sensitive Polymersomes 212 pH-Sensitive Polymer–Inorganic Hybrid Nanoparticles 214 pH-Sensitive Dendrimers 214 pH-Sensitive Polymers for Bioimaging 215 Conclusions 216 References 216

9

Charge-Reversal Polymers for Biodelivery 223 Bo Zhang, Kai Wang, Jingxing Si, Meihua Sui, and Youqing Shen

9.1 9.2

Applications of Cationic Polymers in Biodelivery 223 Barriers for Cationic Polymers in In vitro and In vivo Applications 224 Characteristic pH Gradients in Tumor Interstitium and Endo/Lysosomes 225 Chemistry of Charge-Reversal Polymers Based on Acid-Labile Amides 226 pHe-Triggered Charge-Reversal 228 pHL -Triggered Charge-Reversal 229 Applications of Charge-Reversal Polymers in Biodelivery Systems 230 Charge-Reversal in Cancer Drug Delivery 230 Charge-Reversal in Gene Delivery 232 Charge-Reversal in Protein Delivery 235 Charge-Reversal Incorporated with Inorganic Materials 236 Perspectives 237 References 237

9.3 9.4 9.4.1 9.4.2 9.5 9.5.1 9.5.2 9.5.3 9.5.4 9.6

10

Phenylboronic Acid-Containing Glucose-Responsive Polymer Materials: Synthesis and Applications in Drug Delivery 243 Rujiang Ma and Linqi Shi

10.1 10.2

Introduction 243 PBA-Containing Polymers Operating Under Physiological Conditions 244 Chemically Crosslinked PBA-Based Gels 247

10.3

XI

XII

Contents

10.4 10.5 10.6

Self-Assembled PBA-Based Polymer Micelles 253 Self-Assembled PBA-Based Polymersomes 266 Perspectives 271 References 272

11

Extracellular pH-Activated Nanocarriers for Enhanced Drug Delivery to Tumors 277 You-Yong Yuan, Cheng-Qiong Mao, Jin-Zhi Du, Xian-Zhu Yang, and Jun Wang

11.1 11.2 11.3 11.4 11.5 11.6 11.6.1

Introduction 277 Passive and Active Tumor Targeting 278 Targeting the Extracellular pH (pHe ) in Tumors 279 Extracellular pH-Induced Drug Delivery to Tumors 280 Ligand Exposure by a Shielding/Deshielding Method 281 Surface Charge Reversing Nanoparticles 283 Enhanced Cellular Uptake by Surface Charge Reversing Nanoparticles 283 Overcoming MDR by Surface Charge Reversing Nanoparticles Enhanced Delivery of siRNA by Surface-Charge Reversing Nanoparticles 295 Conclusion 300 References 300

11.6.2 11.6.3 11.7

305

12

Stimulation-Sensitive Drug Delivery Systems Xintao Shuai and Du Cheng

12.1 12.2 12.2.1 12.2.2 12.2.3 12.2.4 12.3 12.4 12.4.1 12.4.2 12.5 12.6

Introduction 305 pH-Sensitive Delivery Systems 306 pH-Sensitive Micellar Delivery Systems 306 pH-Sensitive Polymer–Drug Conjugates 307 pH-Sensitive Dendrimers 308 pH-Sensitive Liposomes 310 Thermo-Sensitive Delivery Systems 311 Biomolecule-Sensitive Delivery Systems 314 Enzyme-Sensitive Nanocarriers 315 Reduction–Responsive Conjugates 316 Other Environmentally Sensitive Nanocarriers 318 Outlook 319 References 320 Index 331

287

XIII

List of Contributors Huabing Chen

Dan Ding

Soochow University College of Pharmaceutical Science Suzhou 215123 China

Nanjing University Laboratory of Mesoscopic Chemistry and Department of Polymer Science & Engineering College of Chemistry & Chemical Engineering 22 Hankou Road Nanjing China

Xuesi Chen

Chinese Academy of Sciences Changchun Institute of Applied Chemistry 5625 Renmin Street Changchun 130022 China Du Cheng

Sun Yat-sen University PCFM Lab of Ministry of Education School of Chemistry and Chemical Engineering Xingangxi Road 135 Guangzhou 510275 China

Jin-Zhi Du

School of Life Sciences University of Science and Technology of China Hefei Anhui 230027 China Guang Hui Gao

Sungkyunkwan University Theranostic Macromolecular Research Center School of Chemical Engineering Republic of Korea

XIV

List of Contributors

Zhishen Ge

Xiqun Jiang

University of Science and Technology of China CAS Key Laboratory of Soft Matter Chemistry Department of Polymer Science and Engineering Hefei 230026 China

Nanjing University Laboratory of Mesoscopic Chemistry and Department of Polymer Science & Engineering College of Chemistry & Chemical Engineering 22 Hankou Road Nanjing China

Pilar Rivera Gil

Centre Tecnològic de la Química de Catalunya (CTQC) and Dpt. Physical and Inorganic Chemistry Dpt. Universitat Rovira i Virgili C. Marcel.li Domingo s/n Edif. N5 43007 Tarragona Spain

Kazunori Kataoka

Zhongwei Gu

Jindˇrich Kopeˇcek

Sichuan University National Engineering Research Center for Biomaterials Chengdu 610064 China

University of Utah Department of Pharmaceutics and Pharmaceutical Chemistry 20 S, 2030 E Salt Lake City Utah 84112 USA

Xiuwen Guan

Chinese Academy of Sciences Changchun Institute of Applied Chemistry 5625 Renmin Street Changchun 130022 China

The University of Tokyo Department of Materials Engineering Graduate School of Engineering 7-3-1 Hongo Bunkyo-ku Tokyo 8656 Japan

and University of Utah Department of Bioengineering Salt Lake City Utah 84112 USA

List of Contributors

Ick Chan Kwon

Moritz Nazarenus

Korea Institute of Science and Technology Biomedical Research Center Republic of Korea

Philipps University of Marburg Department of Biophotonics Institute of Physics Renthof 7 35037 Marburg Germany

Doo Sung Lee

Sungkyunkwan University Theranostic Macromolecular Research Center School of Chemical Engineering Republic of Korea Yi Li

Sungkyunkwan University Theranostic Macromolecular Research Center School of Chemical Engineering Republic of Korea Kui Luo

Sichuan University National Engineering Research Center for Biomaterials Chengdu 610064 China Rujiang Ma

Key Laboratory of Functional Polymer Materials Ministry of Education and Institute of Polymer Chemistry Nankai University Tianjin 300071 China Cheng-Qiong Mao

School of Life Sciences University of Science and Technology of China Hefei Anhui 230027 China

Wolfgang J. Parak

Philipps University of Marburg Department of Biophotonics Institute of Physics Renthof 7 35037 Marburg Germany Kinam Park

Purdue University Departments of Biomedical Engineering and Pharmaceutics 206 S. Martin Jischke Drive West Lafayette Indiana, 47907 USA Youqing Shen

Zhejiang University Center for Bionanoengineering Department of Chemical and Biological Engineering 38 Zheda ST Hangzhou 310027 China Linqi Shi

Key Laboratory of Functional Polymer Materials Ministry of Education and Institute of Polymer Chemistry Nankai University Tianjin 300071 China

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XVI

List of Contributors

Xintao Shuai

Kai Wang

Sun Yat-sen University PCFM Lab of Ministry of Education School of Chemistry and Chemical Engineering Xingangxi Road 135 Guangzhou 510275 China

Zhejiang University Center for Bionanoengineering Department of Chemical and Biological Engineering 38 Zheda ST Hangzhou 310027 China Wenping Wang

Jingxing Si

Zhejiang University Center for Bionanoengineering Department of Chemical and Biological Engineering 38 Zheda ST Hangzhou 310027 China Meihua Sui

Zhejiang University Center for Bionanoengineering Department of Chemical and Biological Engineering 38 Zheda ST Hangzhou 310027 China

Ningxia Medical University Department of Pharmaceutics, School of Pharmacy 1160 Shengli Street Yinchuan Ningxia, 750004 China and Purdue University Departments of Biomedical Engineering and Pharmaceutics 206 S. Martin Jischke Drive West Lafayette Indiana, 47907 USA Jiyuan Yang

Chinese Academy of Sciences Changchun Institute of Applied Chemistry 5625 Renmin Street Changchun 130022 China

University of Utah Department of Pharmaceutics and Pharmaceutical Chemistry 20 S, 2030 E Salt Lake City Utah 84112 USA

Jun Wang

Xian-Zhu Yang

School of Life Sciences University of Science and Technology of China Hefei Anhui 230027 China

School of Life Sciences University of Science and Technology of China Hefei Anhui 230027 China

Huayu Tian

List of Contributors

You-Yong Yuan

Bo Zhang

School of Life Sciences University of Science and Technology of China Hefei Anhui 230027 China

Zhejiang University Center for Bionanoengineering Department of Chemical and Biological Engineering 38 Zheda ST Hangzhou 310027 China and University of Wyoming Department of Chemical and Petroleum Engineering 1000 E. Univ. Ave., Laramie WY 82071 USA

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XIX

Preface The rapid development of biomedical materials science and engineering and advances in modern medicine, pharmacy, biology, and engineering have made the great promise of advanced drug delivery systems (DDS) increasingly clear. These DDS systems are made of carrier materials and drugs, including peptides, proteins, antigens, and nucleic acid drugs, allowing the controlled release of active agents. The advanced DDS systems would not only be a revolutionary change from the traditional mode of drug delivery but may also facilitate the development of currently infeasible approaches to the treatment of cancer, cardiovascular disease, AIDS, congenital genetic defects, and other diseases. It has also greatly promoted the development of molecular diagnostic medicine, which may facilitate the early diagnosis and exploration of the pathogenesis and development of disease and the resultant pathological and physiological changes. With the continuous development of new materials and new technologies and urgent clinical needs, the production of carrier materials and the drug controlled-release/delivery system has become an important part of the entire pharmaceutical industry, and is growing to be the most promising sunrise industry, bringing tremendous and far-reaching impact on the global pharmaceutical industry. The forces driving its rapid growth are as follows: First, the research and development of carrier materials capable of restoring and improving the body’s physiological functions and delivery systems capable of releasing them in a controlled manner are the main direction of contemporary biomedical material research. Second, diseases that are currently difficult to cure may become treatable or easier to detect early. This may involve a significant reduction in health care costs. Third, drug delivery systems capable of controlled release have several advantages over current systems. They tend to last longer, deliver their payload more efficiently in terms of time, and are less toxic to the patient. This also improves the bioavailability of drugs, especially peptide, protein, and nucleic acid drugs. Fourth, these systems are conducive to the development of new drugs. They can reduce costs, shorten development cycles, and provide a quicker return on investment than conventional drug development.

XX

Preface

The designs and preparation techniques of carrier materials and micro-systems are the key to the development of the advanced, efficient drug controlledrelease/delivery system and the theranostic micro devices. The cutting-edge areas and future trends of this field mainly include the following aspects. The first aspect is the intelligent/microenvironment-triggered (stimulusresponsive) DDS. Such systems use the material’s ability to detect changes in pH, redox, other chemical signals, temperature, optics, magnetic fields, electrical signals, mechanical signals, other physical signals, and enzymes, receptors, and other biological signals to facilitate responsive drug release. It may be the most promising and the most valuable DDS with the greatest prospect in clinical applications in the t 21st century. The key technology in this field is to improve the sensitivity of the material and the DDS, in order to achieve rapid response in vivo. The second aspect is the targeted DDS itself, which includes passive and active targeting. Passive targeted DDS cannot recognize the target cells, instead relying on the size effect to reach the target site. For example, a drug carrier system that is 20–200 nm in size triggers an enhanced permeability and retention (EPR) effect based on the defects (holes) in the blood vessels of the tumor tissues, thereby causing efficient local accumulation of drugs. Precise control of the particle size, sufficiently narrow size distribution, and increase in the circulation time in vivo are crucial to passive targeting. Active targeted DDS introduces a targeting group capable of recognizing the target tissue, target cells, or even target molecules. The ability to recognize the cells or molecules of a specific tissue or organ is highly relevant to material preparation. The advantages of this type of DDS are its high selectivity and reduced side effects to normal tissues and organs, which would be very valuable in the treatment of common multiple malignancies. The third aspect is self-regulated DDS. The system must mimic the complex biochemical processes in vivo and release the drug in accordance with the body’s needs, rather than at a stable, predetermined speed. The release mechanisms involve the microporous or bulk diffusion of the polymer reservoir or the enzymatic degradation on the surface of the polymer matrix. The fourth aspect is time control. To achieve constant speed, zero-order release of the drug is the main direction. The chemical structure, composition, and degradation properties of the material and its capacity for drug penetration and diffusion are key to achieving constant release. Transdermal drug delivery avoids the first-pass effect and gastrointestinal damage because it does not require that the drug pass through the liver. This makes it more practical for patients requiring long-term continuous administration. It is also highly efficient and may reduce the rate of side effects. In addition, the source of the drug, usually a patch, can be removed at any time. The core issues in transdermal drug delivery are producing a polymer film capable of drug transmission, which involves the principles of pharmacodynamics and pharmacokinetics. This book covers bioinspired and polymer nano drug delivery systems. It is published jointly by John Wiley Publishing Company and China Chemical Industry Press. Its goal is to review relevant progress and the future direction

Preface

of development in order to promote the development of this field, the clinical application of nano drug delivery systems, and improvements in medicine. To this end, internationally renowned experts in this field co-authored the book with different styles and from different perspectives. The book consists of 12 chapters, mainly including backbone-degradable high molecular weight (second-generation) water-soluble polymer-anticancer drug conjugates, and a new paradigm of drug-free macromolecular therapeutics; nano-targeting drug delivery systems based on peptide dendrimers; nano composite colloidal systems with multiple functions including targeting, stimulus-responsive controlled release, and sensing; multifunctional polymeric micelles that can break various physiological barriers for targeted delivery; biomimetic polymer used for in vivo drug delivery; nanoparticle-based proteins such as gelatin, human serum albumin, collagen, silk protein, casein protein, and elastin-like polypeptides; polymer carriers for gene delivery system and their functional modifications; pHsensitive nano drug delivery systems for targeted cancer therapy and biological imaging; drug/gene nano delivery system capable of charge flipping; nano drug carrier systems using phenylboronic acid-based glucose-responsive polymeric materials and their gels, micelles, and vesicles; nano delivery systems for cancer treatment co-carrying drugs and siRNA and activated by extracellular pH; and pH, heat, and biological molecule or light-sensitive stimuli-responsive DDS. This book covers the design principles, research and development technologies, and application prospects of advanced and efficient polymer nano DDS, and rather comprehensively discusses the outstanding progresses and future trends of the field. This forward-looking, novel treatise on polymer nano DDS is unique and extremely valuable, and I believe that it will be welcomed by readers. In particular, young scholars interested in the field of polymer nano drug delivery systems will learn a great deal from this book, and may be inspired to carry on their own research. I would like to thank all the editors and contributing authors for their time and expertise. I am especially grateful to Dr. Gang Wu from the China Chemical Industry Press and Dr. Esakki Rahini from the John Wiley Publishing Company for their contributions to the publishing of this book. I thank Professor Bin He, Professor Yao Wu, and everyone else who have helped. I would also like to sincerely thank Professor Renxi Zhuo of the Chinese Academy of Science, a pioneer and esteemed leader in the field of polymer chemistry and drug delivery systems, for writing the preface for this book and for his guidance and support. Professor Zhongwei Gu Chief Scientist of Biomaterials, National 973 Program IUS-BSE Fellow Sichuan University Chengdu, China

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1

1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics Jiyuan Yang and Jindˇrich Kopeˇcek

1.1 Introduction

To address the lack of specificity of low molecular weight drugs for malignant cells, the concept of targeted polymer–drug conjugates was developed in the 1970s. The major rationale for the use of water-soluble polymers as carriers of anticancer drugs is based on the mechanism of cell entry [1–3]. Whereas the majority of low molecular weight drugs enter the cell interior by diffusion through the plasma membrane, the entry of macromolecules is restricted to endocytosis [4]. Macromolecules captured by this mechanism are channeled to the lysosomal compartment of the cell. In addition, moieties that complement cell surface receptors or antigens of a subset of cells may be incorporated into the macromolecular structure and render the conjugate biorecognizable [5–9]. There are numerous reviews, which summarize the rationale, design, synthesis, evaluation, and development of macromolecular therapeutics [5–16]. In this chapter, polymeric carriers based on copolymers of N-(2-hydroxypropyl)methacrylamide (HPMA) are used as an example. However, the conclusions can be considered generally applicable to water-soluble carriers with other chemical structures. When using nondegradable water-soluble polymer drug carriers, their molecular weight needs to be below the renal threshold to safeguard biocompatibility. Unfortunately, this leads to relatively short intravascular half-life and the accumulation of the conjugates in solid tumors (because of enhanced permeability and retention (EPR) effect) is suboptimal [17]. Consequently, new designs for macromolecular therapeutics are needed and are discussed. Following a discussion of general design principles of water-soluble polymer–drug conjugates – this chapter focuses on two recent designs: (i) New, second generation anticancer nanomedicines based on biodegradable, high molecular weight HPMA copolymer – drug carriers containing enzymatically degradable bonds in the main chain (polymer backbone). (ii) New paradigm in drug delivery – drug-free macromolecular therapeutics. This approach is based on the biorecognition of complementary motifs at cell surface and crosslinking Bioinspired and Biomimetic Polymer Systems for Drug and Gene Delivery, First Edition. Edited by Zhongwei Gu. © 2015 by Chemical Industry Press. All rights reserved. Published 2015 by Wiley-VCH Verlag GmbH & Co. KGaA.

2

1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics

of receptors with concomitant initiation of apoptosis; no low molecular weight drug is involved.

1.2 Water-Soluble Polymers as Carriers of Anticancer Drugs

The conjugation of drugs to synthetic and natural macromolecules was initiated about 60 years ago – for reviews of early work see Refs. [2, 18]. Jatzkewitz [19] used a dipeptide (glycylleucine) spacer to attach a drug (mescaline) to polyvinylpyrrolidone in the early 1950s and Ushakov’s [20–22] group in Leningrad (now St. Petersburg) synthesized conjugates of polyvinylpyrrolidone with various antibiotics in the 1960s and 1970s. Mathé et al. [23] pioneered the conjugation of drugs to immunoglobulins, setting the stage for targeted delivery. De Duve [1] (who received the Nobel Prize in 1974) discovered that many enzymes are localized in the lysosomal compartment of the cell and the lysosomotropism of macromolecules, which are an important phenomena for the design of polymer–drug conjugates. Finally, Ringsdorf [24] analyzed the research results of the field and presented a clear concept of the use of polymers as targetable drug carriers. 1.2.1 First Generation Conjugates – Design, Synthesis, and Activity

As mentioned in Section 1.1, the design of first generation conjugates was based on the lysosomotropism of macromolecules, on the binding of drugs to the carrier via an attachment/release point that is stable in the blood stream but susceptible to catalyzed hydrolysis in the lysosomal compartment of the cell [25, 26], and on optional attachment of a targeting moiety [5, 6]. Numerous conjugates were evaluated and several proceeded into clinical trials [27–35]. Clinical trials with polymer–drug conjugates demonstrated a decrease of adverse effects when compared to free drugs, but the increase of efficacy (when compared to unbound drugs) was considerably smaller than in animal models. The reasons were recently analyzed and new directions in macromolecular therapeutics research proposed [10]. 1.2.2 Analysis of Design Factors That Need Attention

There are numerous design factors that could speed up the translation of basic research into the clinics [10]. They are briefly discussed in the following sections: 1.2.2.1 Design of Conjugates for the Treatment of Noncancerous Diseases

HPMA copolymer conjugates with the well-established bone anabolic agent (prostaglandin E1 ; PGE1 ) are being developed for the treatment of osteoporosis and other musculoskeletal diseases. The biorecognition of the conjugates by

1.2

Water-Soluble Polymers as Carriers of Anticancer Drugs

the skeleton is mediated by an octapeptide of D-aspartic acid (D-Asp8 ) or a bisphosphonate, alendronate (ALN) [36, 37]. The same approach is applicable for the treatment of bone cancer. 1.2.2.2 Combination Therapy Using Polymer-Bound Therapeutics

Combination therapy using polymer-bound therapeutics has been studied actively [5, 38]. The first combination therapy using polymer-bound drugs used a mixture of HPMA copolymer–DOX (doxorubicin) conjugate and HPMA copolymer – chlorin e6 conjugate [38]. On two cancer models, Neuro 2A neuroblastoma [38] and human ovarian carcinoma OVCAR-3 xenografts in nude mice [39–41] it was shown that combination therapy produced cures that could not be obtained with either chemotherapy or photodynamic therapy alone (Figure 1.1). Incorporation of anti-CD47 antibodies [42] or Fab′ fragments [43, 44] to these conjugates further increased the therapeutic efficacy. From the synthetic and scale-up point of view it is preferable to use a mixture of two conjugates, each containing one drug. However, Vicent et al. [45] have shown that for some drug combinations binding two drugs to the same macromolecule results in higher efficacy when compared to a mixture of two polymer drugs. Recently, a new therapeutic strategy for bone neoplasms using combined targeted polymer-bound angiogenesis inhibitors (two per macromolecule: ALN and antiangiogenic TNP-470) was developed. The bispecific

% Tumor volume

1500 1300

Control

1100

P-Mce6 P-DOX

900

P-DOX+P-Mce6

700 500

300 100 −100

0

5

10

15 Days

Figure 1.1 Combination of chemotherapy and photodynamic therapy of OVCAR3 tumors heterotransplanted in nude mice treated with the HPMA copolymerbound anticancer drugs, doxorubicin (DOX), and chlorin e6 (Mce6 ): control (vehicle); HPMA copolymer–Mce6 conjugate

20

25

30

(P-Mce6 ; 1.5 mg kg−1 Mce6 equivalent) with light; HPMA copolymer–DOX conjugate (P-DOX; 2.2 mg kg−1 DOX equivalent); combination therapy P-DOX + P-Mce6 with light (2.2 mg kg−1 DOX equivalent plus 1.5 mg kg−1 Mce6 equivalent). Bars, SE. (Adapted from Ref. [39].)

3

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1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics

HPMA copolymer ALN-TNP-470 is the first antiangiogenic conjugate that targets both the tumor epithelial and endothelial compartments, warranting its use on angiogenesis-dependent calcified neoplasms such as osteosarcomas and bone metastases [46, 47]. 1.2.2.3 New Targeting Strategies

New targeting strategies involve identification of targeting moieties that can enhance the efficacy of macromolecular therapeutics. One example is the selection of peptides using combinatorial approaches, phage display [48], and the chemical combinatorial technique, one-bead one-compound (OBOC) method [49]. However, the first selection usually produces peptides with a low binding constant and the use of a second, directed library is needed. This is time consuming, but may produce an optimization of the structure and enhancement of the binding constant by several orders of magnitude [50]. Another important aspect for the future development of anticancer nanomedicines is the targeting of cancer stem cells (CSCs) [8, 9, 51]. Cancer cells are biologically and functionally heterogeneous, in terms of phenotype, proliferation, tumorigenesis, invasiveness, and so on. Noticeably, cancer cells are present in various differentiation statuses, with relatively undifferentiated CSCs maintaining the hierarchical organization of the tumor mass, similar to the role of normal stem cells (NSCs) in healthy tissues [52, 53]. Moreover, the CSC theory suggests that the often-observed treatment failures are largely because of the failure of conventional cytotoxic anticancer therapies to eliminate CSCs. Therefore, targeting CSCs or in combination with traditional anticancer therapeutics represents a promising strategy to improve cancer patient survival [51]. Aiming to improve the outcome of prostate cancer treatments by targeting CSCs, we designed a CSC-specific nanomedicine. Cyclopamine, a hedgehog pathway inhibitor, was attached to the end of GFLG (glycylphenylalanylleucylglycyl) biodegradable tetrapeptide side chains of HPMA copolymer. We evaluated the CSC inhibitory effects of the HPMA copolymer-cyclopamine conjugate in an in vitro prostate cancer epithelial cell model using cells derived from a prostate cancer patient that were immortalized by transcription of human telomerase reverse transcriptase (RC-92a/hTERT), [54]. RC-92a/hTERT cells were chosen as the CD133+/integrin α2β1hi /CD44+ putative prostate CSCs within the whole cell line could be enriched to 5%, higher than that reported on primary prostate cancer cells or other established prostate cancer cell lines. Cell surface marker expression analysis and cytotoxicity studies following drug and conjugate treatments on RC-92a/hTERT cells supported the anti-CSC efficacy of the designed macromolecular therapeutics. The HPMA copolymer–cyclopamine conjugate, like free cyclopamine, showed selective inhibitory effect on prostate CSCs when compared with bulk cancer cells in the in vitro prostate cancer model. In contrast, docetaxel, a traditional chemotherapeutic agent for prostate cancer, showed preferential cytotoxicity to bulk cancer cells. These results suggest the treatment potential of a combination of macromolecular therapeutics targeting both bulk tumor cells and CSCs [51].

1.2

Water-Soluble Polymers as Carriers of Anticancer Drugs

1.2.2.4 Relationship between Detailed Structure of the Conjugates and Their Properties

Current research in drug delivery is not sufficiently focused on the analysis of the interplay of individual factors in multifunctional conjugates on the final properties. For example, an increase in the number of hydrophobic targeting peptides per macromolecule leads to enhanced avidity of the conjugate and better targetability. However, during internalization, the conformational changes of the macromolecule may lead to the association of side chains terminated in drug with concomitant decrease in the drug release rate [55, 56]. Combination of characterization techniques that include analysis of the conformation of the polymer conjugate by, for example, FRET (fluorescence resonance energy transfer) needs to be undertaken to optimize the structure of the conjugate. Increased amount of hydrophobic peptide side chain resulted in a more compact conformation of the polymer coil (Figure 1.2). Unsurprisingly, the rate of enzymatically catalyzed release of DOX from the compact coil decreased [55].

Random coil

CH2

Tryptophan

Dansyl

NH

N

Energy donor

Energy acceptor

CH2

NH

N

YILIH

RN

N

IHR

YIL

(b) Compact coil

FRET

Fluorescence intensity

(a)

(c)

Figure 1.2 Evaluation of conformational changes in HPMA copolymer conjugates caused by intramolecular association of hydrophobic (YILIHRN) side chains. Fluorescence (Förster) resonance energy transfer (FRET) measurements using covalently bound tryptophan as donor and dansyl as acceptor could distinguish conformation of HPMA

P-Trp-Dans P-Trp-Dans-pep 1.8% P-Trp-Dans-pep 4.3%

350 300 250 200 150

Blue shift

100 50 0 320

370

420

470

520

Wavelength copolymer containing donor and acceptor only (P-Trp-Dans), HPMA copolymer containing 1.8 mol% of YILIHRN side chains (PTrp-Dans-pep 1.8%) and HPMA copolymer containing 4.3 mol% of YILIHRN side chains (P-Trp-Dans-pep 4.3%). (Adapted from Ref. [55].)

5

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1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics

1.2.2.5 Impact of Binding a Drug to a Polymer on the Mechanism of Action

Free and polymer-bound drugs may activate different signaling pathways because of different internalization mechanisms. This hypothesis is based on the fact that low molecular weight drugs interact with recognition moieties located at cell surface; biorecognition may result in the initiation of a signaling pathway. In contrast, water-soluble polymer-bound drugs are internalized in membrane-limited organelles and thus restricted in the interactions with surface structures. They are released from the carrier in secondary lysosomes and usually enter the cytoplasm in the perinuclear region. Consequently, they will interact with different molecules and could initiate a different signaling pathway. Examples are in the literature on the differences in the mechanism of action of free and HPMA polymer-bound drugs, including DOX [57, 58] and geldanamycin (GA) [59]. For example, the gene expression profiles of human ovarian carcinoma A2780 cells were examined after exposure to free 17-(3-aminopropylamino)17-demethoxygeldanamycin (AP-GA) and HPMA copolymer-bound AP-GA (P(AP-GA)) [59]. As P(AP-GA)-treated cells may exhibit delayed responses, longer exposure times (2 × IC50 dose for 6 and 12 h) were examined (Figure 1.3). The hierarchical clustering of the expression ratios of the selected 68 genes indicated considerable similarities in the gene expression profiles after APGA and P(AP-GA) treatments (Figure 1.3a). On the other hand, unlike the a

b

c

d

a. P(AP-GA) 12 h HSC71

HSP70

HSP90 HSP27 b. P(AP-GA) 6 h HSC71

HSP70

HSP90 HSP27 Down-regulation (a)

c. AP-GA 12 h HSC71

HSP70

HSP90 HSP27 d. AP-GA 6 h HSC71

HSP70

HSP90 HSP27 Up-regulation

Zero

(b)

Figure 1.3 Gene expression in A2780 ovarian carcinoma cells exposed to free 17-(3-aminopropylamino)-17demethoxygeldanamycin (AP-GA) and HPMA copolymer bound AP-GA (P(AP-GA)) at 2 × IC50 concentration for 6 and 12 h. (a)

Hierarchical cluster analysis of the expression of selected 68 genes in the Atlas human 1.2 cDNA expression array. (b) Gene expression of cell stress response-related proteins. (Adapted from Ref. [59].)

1.2

Water-Soluble Polymers as Carriers of Anticancer Drugs

AP-GA treatment, P(AP-GA) treatment induced little expression in stress response-related genes even after 12 h (Figure 1.3b). As GA-treated cells exhibited little expression in stress response-related genes, the elevated expression of stress response-related genes after exposure of cells to AP-GA may not be directly related to cell death mechanism induced by HSP90 inhibition. Therefore, it is probable that P(AP-GA) may suppress the expression of stress response-related genes activated by AP-GA following differences in its internalization mechanism, subcellular localization, and intracellular concentration gradients. Thus the results suggest that conjugation of AP-GA to HPMA copolymer may be able to modulate the cell stress responses induced by AP-GA following differences in its internalization mechanism, subcellular localization, and intracellular concentration gradients [59]. 1.2.2.6 Mechanism of Internalization and Subcellular Trafficking

Macromolecular therapeutics cannot cross the phospholipid bilayer by diffusion; they enter cells by endocytic pathways [4]. Most common classification schemes of endocytosis are based on protein machinery that facilitates the process, such as clathrin-mediated endocytosis and clathrin-independent endocytosis [60–63]. Clathrin-independent endocytosis is further categorized as caveolae-mediated endocytosis and clathrin- and caveolin-independent endocytosis [60, 62] or dynamin-dependent and dynamin-independent endocytosis [62, 63]. In addition, macropinocytosis is a distinct pathway of pinocytosis [64]. The relationship between the detailed structure of the polymer–drug conjugate and its mechanism of internalization is important information, which provides feedback for the optimization of the conjugate structure. Recently, research has been focusing on the identification of different routes of cell entry with the aim to deliver drugs into subcellular compartments different from lysosomes. As the activity of many drugs depends on their subcellular location, manipulation of the subcellular fate of macromolecular therapeutics may result in more effective conjugates. Approaches that seem to be effective are nuclear delivery of drugs mediated by steroid hormone receptors that shuttle between the cytoplasm and the nucleus [5, 65] and mitochondrial targeting mediated by delocalized hydrophobic cations [66–69]. Particularly, the experiments of Murphy et al. [68, 69] used terminally functionalized triphenylphosphonium (TPP) to target antisense peptide nucleic acid (PNA) into the mitochondria of isolated organelles and whole intact cells in vitro. Attachment of TPP to HPMA copolymer resulted in enhanced mitochondrial localization following microinjection and incubation experiments with ovarian carcinoma cells [66, 67]. Recently, Torchilin et al. used the same concept with TPP modified dendrimers [70] and liposomes [71]. 1.2.2.7 Relationship between the Molecular Weight of the Carrier and the Efficacy of the Conjugate

High molecular weight (long-circulating) polymer conjugates accumulate efficiently in tumor tissue because of the EPR effect [72]. Experimental data have shown that the higher the molecular weight of the conjugate, the higher the

7

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1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics

accumulation in the tumor tissue with concomitant increase in therapeutic efficacy [73]. However, the renal threshold limits the molecular weight of the first generation of polymeric carriers to below ∼50 kDa; this lowers the retention time of the conjugate in the circulation with concomitant decrease in pharmaceutical efficiency. Higher molecular weight drug carriers with a nondegradable backbone deposit and accumulate in various organs, impairing biocompatibility. To this end we designed the new second generation anticancer nanomedicines based on high molecular weight HPMA copolymer–drug carriers containing enzymatically degradable bonds in the main chain (polymer backbone) [74–76]. These multisegment block copolymers are synthesized by reversible addition-fragmentation chain transfer (RAFT) polymerization followed by click (alkyne–azide and/or thiol–ene) reactions (see following paragraph). 1.2.3 Design of Second Generation Conjugates – Long-Circulating and Backbone Degradable

HPMA copolymer-anticancer drug conjugates have been evaluated in clinical trials, but no product has entered clinical use yet. Questions have been raised whether HPMA conjugates “have a future as clinically useful nanomedicines?” [14]. In the following sections, we summarize our efforts in the design, development, and evaluation of second generation backbone degradable HPMA copolymer–anticancer drug conjugates. The in vitro and in vivo data suggest a great potential of these new nanomedicines. Backbone degradable HPMA copolymer–drug conjugates were designed and synthesized by combination of RAFT copolymerization with either alkyne–azide or thiol–ene click chemistry [74–76]. An example, the synthesis of HPMA copolymer–gemcitabine (GEM) conjugates, is shown in Figure 1.4. Compared with previously evaluated HPMA, copolymer conjugates entered early clinical trials but were discontinued; the hallmark of the new second generation conjugates is the biodegradability of the HPMA linear backbone because of multiblock structure composed of alternating HPMA copolymer blocks and enzyme-cleavable oligopeptide segments. This permits to use high molecular weight long-circulating conjugates without impairing biocompatibility. Another remarkable synthetic feature is the utilization of polymerizable drug derivatives and controlled polymerization chemistry rather than attachment of drugs via polymer-analogous reactions. The drug content and average Mw can be tailored by the variation of feed ratio and polymerization conditions. The process results in the narrow distribution of molecular weights and minimal heterogeneity in chemical composition of the conjugates. In addition, the amount of free drug in the conjugates is minimized. 1.2.3.1 RAFT Copolymerization for the Synthesis of Conjugates

The advances in controlled (“living”) radical polymerization contributed to the improved designs of nanomedicines by permitting the synthesis of well-defined

1.2

Water-Soluble Polymers as Carriers of Anticancer Drugs

O 1.

NH

S

O O NH

+

CN

H N

S

NH O

O

2.

NH

CN

N

N H

O

NC

N

N3

N3

O F

H N

GFLG

H N

O

OH F

O

CuBr/L-ascorbic acid, DMF, rt

MA-GFLG-GEM

Tetramer G4

GEM

α,ω-Dialkyne telechelic HPMA copolymer G0 , or 2P-GEM

N O N

Multiple segment

GEM NH O

70 °C, 3 h, DMSO

HN

HO

G F L G

V501, 70 °C, H2O/DMSO, 16 h

O

HPMA

G F L G

S O

Peptide2CTA

(2-arm- CTA with GFLG sequence)

NH

HO

GFLG

S

NC

H

GFLG N

Lys

GFLG

O

Fractionation

Trimer

Dimer

G3

G2

GFLG

GFLG

G0

(a)

GFLG G F L G

G F L G

G F L G

G F L G

GEM

GEM

GEM

G2

GEM

G0

After degradation

G3

G0 (2P-GEM)

(b)

After click/fractionation

G2

G3

Mw 57 kDa Mw/Mn1.10

2P-GEM Mw 110 kDa Mw/Mn1.12

G4

Mw, kDa

110

213

314

427

Mw/Mn

1.12

1.06

1.17

1.06

Gem%

5.6

5.5

4.8

−

0

(c)

5

10

15

20

Elution volume (ml)

25

5

(d)

10

15

20

25

Elution volume (ml)

Figure 1.4 HPMA copolymer–gemcitabine conjugates. (a) (i) Synthesis of backbone degradable multiblock conjugates via three steps: (1) RAFT copolymerization using peptide2CTA as RAFT agent; (ii) chain end modification to obtain clickable telechelic diblock drug conjugates; and (iii) chain extension by alkyne–azide click reaction. (2) Synthesis of first generation conjugate (P-GEM). (b) Characterization of narrow molecular weight fractions of HPMA copolymer–gemcitabine conjugates: G0, diblock HPMA copolymer conjugate prepared in one step using the new CTA containing a GFLG degradable sequence; G2, tetrablock (dimer of dimer) isolated by fractionation; G3, hexablock (trimer of diblock) isolated by fractionation. (c) Size exclusion chromatograms of G0 and G2 and G3 fractions. (d) Following incubation with papain, the molecular weight of 2P-GEM decreased to half of the original value. (Adapted from Ref. [77].)

9

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1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics

bioconjugates that possess narrow distribution of molecular weights and a low degree of compositional (chemical) heterogeneity. The use of RAFT polymerization with dithiobenzoate or trithiocarbonate chain transfer agents (CTAs) [78] is an excellent approach for water-soluble polymer–anticancer drug conjugates, including the (co)polymerization of HPMA [74–76, 79]. The use of CTAs that contain an enzymatically degradable sequence in their structure results in polymer chains with inserted degradation point. Yang et al. synthesized an enzyme-sensitive, alkyne functionalized CTA, N α -(4-pentynoyl)-N δ -(4- cyano-4(phenylcarbonothioylthio)pentanoylglycylphenylalanylleucylglycyl)-lysine(CTAGFLG-alkyne), and used it in the synthesis of polyHPMA that contained an alkyne attachment point connected via an enzymatically degradable bond (Figure 1.5). Postpolymerization modification with 4,4′ -azobis(azidopropyl 4-cyanopentoate) resulted in the formation of heterotelechelic HPMA copolymers containing terminal alkyne and azide groups [74]. Similarly, Pan et al. synthesized an enzyme-sensitive bifunctional CTA, N α , ϵ N -bis(4-cyano-4-(phenylcarbonothioylthio)pentanoylglycylphenylalanylleucylglycyl)lysine (Peptide2CTA). During RAFT polymerization the HPMA monomers incorporated at both dithiobenzoate groups. When the final polymer was incubated with papain – a thiol proteinase with similar specificity as lysosomal proteinases – the molecular weight decreased to half of the original value; this suggests that the monomers inserted into both dithiobenzoate groups of the Peptide2CTA with identical efficiency (Figure 1.4). The postpolymerization aminolysis of the terminal dithiobenzoate moieties results in telechelic α,ω-dithiol-polyHPMA [76]. 1.2.3.2 Click Reactions for Chain Extension into Multiblock Copolymers

The RAFT copolymerization of HPMA with polymerizable derivatives of anticancer drugs and subsequent postpolymerization end-group modification results in heterotelechelic or homotelechelic copolymers that can be chain extended into high molecular weight multiblock backbone degradable copolymers. However, the chain extension is a condensation reaction that widens the distribution of molecular weights. Consequently, the product needs to be fractionated by size exclusion chromatography to create fractions with narrow polydispersity. For example, the heterotelechelic HPMA copolymers described above (Figure 1.5) can be chain extended just by exposing them to a catalyst facilitating the alkyne–azide reaction [74]. Similarly, telechelic α,ω-dialkyne-polyHPMA can be chain extended by a diazide containing an enzymatically degradable sequence [75]. The telechelic α,ω-dithiol-polyHPMA can be chain extended by dimaleic anhydride compounds, for example, bis-[1,13-(3-maleimidopropionyl)amido]4,7,10-trioxatridecane (Bis-MAL-dPEG3) in methanol at room temperature for 24 h [76]. The schematic of chain extension reactions is shown in Figure 1.6. 1.2.3.3 Biological Properties of Long-Circulating Macromolecular Therapeutics

The improved therapeutic efficacy of second generation, backbone degradable HPMA copolymer conjugates has been recently evaluated in several studies.

1.2

O

CN

+ NH

HO

N H

O

NH O

O

H N

S

O NH

O

H N

H N

H N

N H

O

COOH O

O

O

GFLG

Cleavage NH O

HN

HO

N3

Tetrapeptide GFLG

O

HPMA N O N O F

11

1. S

NH

Water-Soluble Polymers as Carriers of Anticancer Drugs

2.

O

NC N3

O

N N

O

CN

O

N3

G F L G

Cleavage GEM

OH F

MA-GFLG-GEM Figure 1.5 Synthesis of clickable heterotelechelic HPMA copolymer–gemcitabine conjugate via RAFT copolymerization using an enzymesensitive (degradable) alkyne functionalized chain transfer agent (CTA), Nα -(4-pentynoyl)-Nδ -(4-cyano-4-(phenylcarbonothioylthio)pentanoylglycylphenylalanylleucylglycyl)-lysine (CTA-GFLG-alkyne). Postpolymerization modification with 4,4′ -azobis(azidopropyl 4-cyanopentoate) resulted in the formation of heterotelechelic HPMA copolymers containing terminal alkyne and azide groups and an enzymatically degradable GFLG sequence sequence. (Adapted from Ref. [74].)

12

1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics CN

S O NH HO

NC

H H N GFLGKGLFG N

S O

(Peptide2CTA)

S

CN N

NH O

Change to thiol via aminolysis

HS

GFLG

GFLG

O

GFLG

O

SH

O H N

N

N3

O

GFLG

H N

N O

O

Thiol-ene reaction

Azide–alkyne cycloaddition

GFLG

O

S

OH

Radical-induced end-group modification

N3

S n/2

NH

HO

N NC

O

O

O

NC

H H N GFLGKGLFG N

NH

O N H

n/2

S

AIBN/MeOH, 50 °C, 20 h

HPMA

CN

S

S

O

GFLG

GFLG

Figure 1.6 Alkyne–azide and thiol–ene chain extension reactions used in the synthesis of backbone degradable, long-circulating HPMA copolymer–drug conjugates. (Adapted from Refs. [74–76].)

Multiblock HPMA copolymer–DOX conjugates were synthesized by RAFT polymerization followed by chain extension via thiol–ene click reaction. The examination of molecular weight dependent antitumor activity toward human ovarian A2780/AD carcinoma xenografts in nude mice revealed enhanced activity of multiblock, second generation higher molecular weight conjugates (Mw = 93; 185; and 349 kDa) when compared to traditional HPMA copolymer–DOX conjugate (Mw = 20 kDa). The examination of body weight changes during treatment indicated the absence of nonspecific adverse effects [80]. In another study, a multiblock backbone degradable HPMA copolymer– paclitaxel conjugate (mP-PTX; Mw = 335 kDa) was synthesized by RAFT copolymerization, followed by chain extension. The in vivo efficiency of free PTX, HPMA copolymer–PTX conjugate with Mw of 48 kDa (P-PTX), and mP-PTX was determined on female nu/nu mice bearing orthotopic A2780 ovarian tumors. Pharmacokinetics study showed that high Mw mP-PTX was cleared more slowly from the blood than commercial PTX formulation and low Mw P-PTX. SPECT/CT imaging and biodistribution studies demonstrated biodegradability as well as elimination of mP-PTX from the body. The tumors in the mP-PTX treated group grew more slowly than those treated with saline, free PTX, and P-PTX (single dose at 20 mg PTX/kg equivalent). Moreover, mice treated with mP-PTX had no obvious ascites and body weight loss. Histological analysis indicated that mP-PTX had no toxicity in liver and spleen, but induced massive cell death in the tumor [81]. Second generation HPMA copolymer–GEM conjugates and combination of HPMA copolymer–GEM and HPMA copolymer–PTX conjugates have

1.2

Water-Soluble Polymers as Carriers of Anticancer Drugs

1600

Dose: 10 mg kg−1 i.v. injection

% Tumor volume

Gemcitabine (free drug)

1400

P-GEM (40 kDa)

1200

2P-GEM (100 kDa) mP-GEM (314 kDa)

1000 800

21-day cycle

600

2 injections followed by 1 week break

400 200 0 0

5

10

15

20

25 30 Days

35

40

45

50

55

Figure 1.7 Molecular weight-dependent in nude mice. The mice received 10 mg kg−1 in vivo antitumor activity of free and HPMA GEM equivalent on days 0, 7, 21, and 28. copolymer-bound gemcitabine (GEM) against Reprinted from Ref. [77] with permission. A2780 human ovarian carcinoma xenografts

shown excellent activity toward human ovarian carcinoma xenografts [77, 82]. The evaluation of HPMA copolymer–GEM conjugates demonstrated the enhanced activity of the diblock (100 kDa) and multiblock (314 kDa) conjugates when compared to the first generation conjugate (40 kDa) (Figure 1.7) [77]. In vivo behavior of a combination of diblock backbone degradable HPMA copolymer–drug conjugates (2P-PTX and 2P-GEM) was investigated using pharmacokinetics, biodistribution, and SPECT/CT imaging studies. In parallel, the antitumor efficacy of combination treatment of 2P-PTX and 2P-GEM was evaluated and compared with free drugs (PTX and GEM) and first generation low Mw conjugates (P-PTX and P-GEM) in nu/nu mice bearing A2780 tumor xenografts. Compared to first generation low Mw HPMA copolymer conjugates, high Mw backbone biodegradable HPMA copolymer carriers significantly prolonged the intravascular half-life of drugs (PTX and GEM) in mice. The biodistribution and SPECT/CT imaging results demonstrated accumulation of conjugates 2P-PTX and 2P-GEM in the tumors and the degradation of new generation conjugates in mice. Notably, the tumors treated with combination of 2P-PTX and 2P-GEM were more effectively repressed, when compared to free drug combination and first generation (low Mw) conjugates combination (Figure 1.8). The histological analysis indicated that the combination treatment had no toxicity in major organs [82]. Furthermore, the backbone degradable HPMA copolymer–GEM conjugate exhibited excellent antitumor activity toward PANC1 pancreatic adenocarcinoma xenografts in nude mice [83].

13

1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics

Tumor model and inoculation (4× 106)/100 μl A2780 cells s.c. injected to 6-8 week old female nude mice Administration: Sequential combination treatment (paclitaxel (PTX) followed by gemcitabine) was given when tumor grew till ~ 50 mm2. The procedure and doses: PTX GEM P-PTX P-GEM 2P-GEM iii. 2P-PTX iv. Control (saline) i. ii.

PTX/PTX conjugates 0

(a)

7

14

21

1

28

7

GEM/GEM conjugates 14 21 28

Days

The dose 5 mg kg−1, i.v. injection, on days 1, 7 and 14 of each 3weeks cycle

One dose 20 mg kg−1, i.v. injection on day 0 of each 3-weeks cycle

7000 PTX + GEM (Free drugs)

6000

P-PTX + P-GEM % Tumor volume

14

5000

2P-PTX + 2P-GEM Saline (control)

4000 3000 2000 1000 0 0

7

14

(b) Figure 1.8 (a, b) Treatment of s.c. A2780 human ovarian carcinoma xenografts by combination therapy of GEM and PTX and their HPMA copolymer conjugates [82]. Characterization of conjugates (Mw in kDa/polydispersity/content of drug in wt%):

21

28

35

Days P-PTX 50/1.08/8.5; 2P-PTX 146/1.41/8.2; P-GEM 32/1.07/7.7; 2P-GEM 89/1.07/7.9. Dosing: 20 mg kg−1 PTX equivalent on day 0 followed by 5 mg kg−1 GEM equivalent on days 1, 7, and 14. (Adapted from Ref. [82].)

1.2.4 Summary of Part 2 and Future Prospects

The development of backbone degradable, linear HPMA copolymers as drug carriers is an important one in the area of water-soluble polymeric drug carriers. It permits to manipulate the intravascular half-life of polymer–drug conjugates in a wide range without impairing biocompatibility. The development of such carriers was possible following the development of new polymerization techniques, namely RAFT [78] and ATRP (atom transfer radical polymerization) [84].

1.3

Drug-Free Macromolecular Therapeutics – A New Paradigm in Drug Delivery

In combination with click reactions a rational design of the second generation conjugates could be achieved. Of particular interest is the development of RAFT CTAs that contain enzymatically degradable sequences in their structure. The possibility to prepare diblock copolymers that contain an enzymatically degradable sequence in one operation is important for the scale-up of the process. A biocompatible polymer carrier with a molecular weight of twice the renal threshold might be sufficient for augmenting the efficacy of water-soluble polymer–anticancer drug conjugates, as experimental results suggest [77, 80–83]. In general, the research in the area of water-soluble polymer–drug conjugates is on good track. The biocompatibility of the polymer carriers as well as the decrease of adverse effects have been proven in clinical trials. The design principles for more efficient conjugates have been identified. Employing modern imaging techniques [85–89] that permit noninvasive monitoring of the fate of conjugates will undoubtedly contribute to a more rational design of polymer therapeutics and theranostics. The major challenge of the field, however, is the translation into clinical application. The FDA approval of a conjugate that will be successful in clinical use will stimulate research and ultimately produce numerous new conjugates.

1.3 Drug-Free Macromolecular Therapeutics – A New Paradigm in Drug Delivery

An exciting new development in the nanomedicine research area is the design of drug-free macromolecular therapeutics. The new paradigm in drug delivery is based on the biorecognition of natural (e.g., peptide) motifs at cell surface, formation of heterodimers (e.g., antiparallel coiled-coils), crosslinking of noninternalizing receptors, and initiation of apoptosis [90–92]. 1.3.1 Biorecognition in Hybrid Polymer Systems

Hybrid polymer systems are composed from at least two distinct classes of macromolecules, for example, synthetic and biological macromolecules. For instance, conjugation of peptide domains to synthetic polymers may produce materials with properties superior to individual components. The peptide domain may insert a level of control over structure resulting from self-assembly at a nanometer scale; the synthetic part may enhance the biocompatibility of the whole system [90]. We have designed hybrid systems composed from hydrophilic HPMA copolymer backbone grafted with two pentaheptad oligopeptide sequences with opposite charge (CCE and CCK). A mixture of equimolar solutions of P-CCK (P is the HPMA copolymer backbone) and P-CCE self-assemble into hydrogels. This process is mediated by the recognition of CCE and CCK peptide grafts – they fold into antiparallel coiled-coils [93, 94].

15

1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics

1.3.2 Coiled-Coils in Biomedical Systems

a,d -- hydrophobic e,g -- charged

abcdefg n

c

g

e d

b

a

f

A coiled-coil trimer

The coiled-coil is one of the basic folding patterns of native proteins. It consists of two or more right-handed α-helices winding together to form (usually) a slightly left-handed superhelix [90, 95–98]. The primary structure of the coiled-coil motif is characterized by a sequence of repeating heptads (motif of seven amino acids) designated as [a, b, c, d, e, f , g]x , in which a and d are usually hydrophobic amino acid residues, while the others are polar. Two helices associate through a hydrophobic interface between a and d, making b, c, and f face outward. Interhelical electrostatic interactions between residues e and g contribute to the stability of the coiled-coil (Figure 1.9). Depending on their detailed structure, α-helices may associate as homodimers or heterodimers in parallel or antiparallel alignments, or form higher order (e.g., tetramer) aggregates [99–101]. Hundreds of native proteins, such as muscle proteins, transcription factors, cytoskeletal proteins, cell and viral surface proteins, tumor suppressors, molecular motors, and many disease- and organ-specific autoantigens, have functional coiled-coil

A coiled-coil dimer

16

f a

d

b e

g

C

c

N

N N

Anti-parallel

C

N

Ubiquitously found in structural proteins, receptors, transcription factors, oncogenes, etc.

C

C Parallel

Close correlation between primary sequence and molecular recognition, oligomerization, and stimuli-sensitive conformational change Figure 1.9 Coiled-coil is a common folding motif. The primary sequence of a typical coiled-coil is composed of seven-residue repeats, designated as heptads. The amino acid residues in a heptad are conventionally denoted as “a, b, c, d, e, f, g.” Hydrophobic residues at positions “a” and “d” form an interhelical hydrophobic core, providing a stabilizing interface between the helices.

Charged residues at positions “e” and “g” form electrostatic interactions, which contribute to coiled-coil stability and mediate specific association among helices. Depending on the primary structure, the helices may form homodimers or heterodimers and associate in a parallel or antiparallel arrangement. (Adapted from Refs. [90, 101].)

1.3

Drug-Free Macromolecular Therapeutics – A New Paradigm in Drug Delivery

domains [102]. A distinctive feature of coiled-coils is the specific spatial recognition, association, and dissociation of helices, making it an ideal model for protein biomaterials in which the higher order structures may be predicted based on the primary sequence. Various functional groups may be exactly positioned into the coiled-coil structure, allowing specific intermolecular interactions to occur. A typical α-helix is right-handed and 3.6 amino acid residues are needed to form a full turn. In a left-handed coiled-coil (composed of right-handed helices), one heptad forms exactly two turns (so-called 7/2 repeat – 7 amino acids per 2 turns). In nature, coiled-coils with different priodicities, for example, 11-residues periodicities, or with insertions of one or more residues into the heptad pattern, can be found (insertions of one residue are called skips, three-residue insertions stammers, and four residue insertions are stutters) [95, 103]. The versatility of the coiled-coil motif, especially the possibility to manipulate its stability and specificity by modifying the primary structure (up to 10−15 M stabilities may be achieved [104]), bodes well for their use in the successful design of new biomaterials. 1.3.3 Coiled-Coil Based Drug-Free Macromolecular Therapeutics: Design, In Vitro, and In Vivo Activity

As mentioned above, we designed two oppositely charged heptapeptad sequences, CCE and CCK, that form antiparallel coiled-coil heterodimers [93]. These sequences served as physical crosslinkers in the self-assembly of HPMA graft copolymers – P-CCK and P-CCE [93, 94] – and in the design of tandem modular proteins [105]. The excellent biorecognition of the peptide domains was an inspiration for the design of new nanomedicines; this created a bridge between the design of biomaterials and the design of nanomedicines. CCK and CCE peptides were engaged in the design of a new CD20+ cell apoptosis induction system, called drug-free macromolecular therapeutics [91, 92]. CD20 is an ideal target for immunotherapies. It is an integral membrane protein [106] that is expressed from pre-B cells to terminally differentiated plasma cells and is present on greater than 90% of B cell malignancies [107, 108]. CD20 is not shed from the cell surface nor is it present in serum under standard physiological conditions. It is a cell cycle regulatory protein [109] that either controls or functions as a store-operated calcium channel. The protein forms dynamic dimers and tetramers [110] constitutively associated with lipid rafts of the cell membrane [111]. Indeed, the biorecognition of CCE/CCK peptide motifs at the cellular surface was able to control apoptosis of CD20+ B cells. Exposure of Raji B cells to an anti-CD20 Fab′ -CCE conjugate decorated the cell surface with CCE (CD20 is a noninternalizing receptor) through antigen–antibody fragment recognition. Further exposure of the decorated cells to CCK-P (grafted with multiple copies of CCK) resulted in the formation of CCE/CCK coiled-coil heterodimers at the cell surface. This second biorecognition induced the crosslinking of CD20 receptors

17

1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics

B cell b. Decorated B-cell with CCE

ic n er atio dim m ro or te il f He d-co ile co

An tig e re n–a ac nt tio ibo n dy

18

Apoptosis

Fab’-CCE

CD20 antigen

CCE random coil random coil CCK

pH 7 CCE/CCK coiled-coils

In

du

P-CCK

B cell

c ti

on

a. Malignant B-cell Figure 1.10 Drug-free macromolecular therapeutics. Cartoon of overall design and possible mechanism of treatment of NHL with conjugates of antiparallel coiledcoil forming peptides, CCE and CCK. Exposure of malignant B cells to anti-CD20 Fab′ fragment – CCE conjugate (Fab′ -CCE) decorates the cells with the CCE peptide by biorecognition of the Fab′ fragment by the noninternalizing CD20 receptor. Further exposure of decorated cells to a HPMA copolymer grafted with several copies of

B cell c. CD20 antigen at cell surface crosslinked complementary CCK peptide(P-CCK) results in the formation of antiparallel coiled-coil at the cell surface, crosslinking of CD20 receptors, and initiation of apoptosis. Inset: Formation of antiparallel coiled-coils by mixing equimolar amount of CCE (maleimideYGG E VSALEKE VSALEKK NSALEKE VSALEKE VSALEK) and CCK (CYGG K VSALKEK VSALKEE VSANKEK VSALKEK VSALKE) peptides. Design of peptides from Ref. [93]; principle of drugfree macromolecular therapeutics. (Adapted from Ref. [91].)

and triggered the apoptosis of Raji B cells in vitro [91] and in a Non-Hodgkin lymphoma animal model in vivo [92]. This is a new concept, where the biological activity of drug-free macromolecular therapeutics is based on the biorecognition of peptide motifs (Figures 1.10 and 1.11). 1.3.4 Potential, Limitations, and Future Prospect of Drug-Free Macromolecular Therapeutics

The design of drug-free macromolecular therapeutics is a truly novel approach. It builds on the design of new self-assembling biomaterials [90, 93, 94, 112, 113] and translates the biorecognition principles to nanomedicine [91, 92, 114]. The first design of this system was tailored for the noninternalizing CD20 receptor and NHL. The preference for noninternalizing receptors is the limitation of the design.

Apoptotic cells (%)

60

0

Drug-Free Macromolecular Therapeutics – A New Paradigm in Drug Delivery

(b)

CCK + CCE

−10

Feb’-CCE

(d)

CCK-P

−20

0

2h

−30 −40 200 210 220 230 240 250 Wavelength (nm)

Caspase 3 assay

4h

1

2

Raji B cells i.v. injection

3

4

5

Mice were monitored for onset of hindlimb paralysis for up to 100 days

Day

Control P = 0.0002 CS P = 0.1189 PS P = 0.0002 P = 0.0001 CM P = 0.1096 PM

1 μM/25 μM

Ctrl 1: Fab’ + CCE + CCK + P Ctrl 2: Fab’-CCE + P Ctrl 3: Fab’ + P-CCK

6

Additional injections of peptide-conjugates (PM and CM)

50 40

19

Treatment with peptide-conjugates (PS, CS, PM, and CM)

1h

100

0.5 μM/25 μM

1 μM/25 μM

30 0.5 μM/25 μM 20 10

7/7

80

Survival (%)

(a)

[θ] × 10−3 (deg.cm2.dmol−1)

1.3

5/7 60 40 20

0 Untreated

Ctrl 1

Ctrl 2

Ctrl 3

Fab’-CCE+P-CCK

Ctrl 1

Ctrl 2

Ctrl 3

Fab’-CCE+P-CCK 1F5+GAM

0 0

(c)

Premixed

Consecutive

20

40

60

80

100

Days after tumor injection

Figure 1.11 Coiled-coil mediated induction of apoptosis in Raji B cells in vitro and in vivo. (a) CD spectra of equimolar mixture of CCE and CCK; (b) biorecognition of Fab′ -CCE (Rhodamine Red labeled) and P-CCK (FITC labeled) on Raji B surface: first column – red channel, second column – green channel, third column – overlay of red and green channels; (c) induction of apoptosis as determined by caspase 3 assay (1F5 + GAM is positive control – exposure to 1F5 Ab followed by goat antimouse secondary Ab; and (d) Therapeutic efficacy of drug-free macromolecular therapeutics against systemically disseminated Raji B cell lymphoma in C.B.-17 SCID mice (7 mice per group). Top panel shows timeline for the in vivo efficacy study. Four million Raji B cells were injected into the tail vein on day 0 to initiate the disseminated disease. The incidence of hind-limb paralysis or survival of mice was monitored until day 100. Five groups of animals were evaluated: untreated controls; consecutive administration of single dose (CS); premixed administration of single dose (PS); consecutive administration of three doses at days 1, 3, and 5 (CM); and premixed administration of three doses at days 1, 3, and 5 (PM). Consecutive administration involved the i.v. injection of 50 μg/20 g Fab′ -CCE first and 1 h later the i.v. administration of 324 μg/20 g P-CCK conjugate; For premixed administration, the two conjugates were mixed together 1 h before injection via the tail vein. Bottom panel shows survival rate of tumor-bearing mice that received above treatments. The curve was presented in a Kaplan–Meier plot with indication of numbers of long-term survivors. (in vitro data adapted from Ref. [91; in vivo data adapted from Ref. [92].)

20

1 Backbone Degradable and Coiled-Coil Based Macromolecular Therapeutics

CD20 is highly expressed on the surface of malignant and normal B cells, but not in stem cells or plasma cells. Thus, drug-free macromolecular therapeutics (employing the “B cell depletion strategy”) can be potentially used to treat B cell-derived hematological neoplastic diseases and autoimmune diseases, without nonreversible impact on normal immune function [115]. Other potential disease targets are (in addition to NHL): chronic lymphocytic leukemia (CLL), rheumatoid arthritis, multiple sclerosis, systemic lupus erythematosus, autoimmune hemolytic anemia, pure red cell aplasia, idiopathic thrombocytopenic purpura, Evans syndrome, vasculitis, bullous skin disorders, type 1 diabetes mellitus, Sjögren’s syndrome, Devic’s disease, and Graves’ ophthalmopathy. All of the above listed diseases have been treated by rituximab anti-CD20 mAb (either approved by FDA or in clinical trials). Importantly, the concept of drug-free macromolecular therapeutics could be expanded by using different components in the design. For example, the Fab′ fragment can be replaced by antigen binding saccharides [116], by peptides selected by phage display [48] or by combinatorial methods [49]. The peptide part can be replaced by complementary oligonucleotides that hybridize at cell surface [114].

1.4 General Summary and Outlook

The advantages of polymer-bound drugs (when compared to low molecular weight drugs) are [5–8]: (i) active uptake by fluid-phase pinocytosis (nontargeted polymer-bound drug) or receptor-mediated endocytosis (targeted polymer-bound drug), (ii) increased passive accumulation of the drug at the tumor site by the EPR effect, (iii) increased active accumulation of the drug at the tumor site by targeting, (iv) long-lasting circulation in the bloodstream, (v) decreased nonspecific toxicity of the conjugated drugs, (vi) potential to overcome multidrug resistance, (vii) decreased immunogenicity of the targeting moiety, (viii) immunoprotecting and immunomobilizing activities, and (ix) modulation of the cell signaling and apoptotic pathways. In addition to preclinical evaluation on animal cancer models, these advantages were recognized in numerous clinical trials of water-soluble polymer–drug conjugates [reviewed in 35]. However, the translation of laboratory research into the clinics has been slow. The approaches to enhance the development and translation have been reviewed here. The field of water-soluble polymer–drug conjugates is at the crossroads. Scientifically, the design principles for bioconjugates are well defined; the challenge is to combine the efficient design of the conjugates with the understanding of the biological features of cancer, including heterogeneity of cancer cells, tumor microenvironment, and metastasis [8, 9]. The progress will occur on several levels, including: (i) continuous progress of our knowledge resulting in the design of bioconjugates with higher activities. Some examples of these strategies were described above, such as the design of conjugates for the treatment of musculoskeletal diseases, combination

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Acknowledgments

We are truly indebted to our coworkers for their support; their contributions are reflected in the references. The research was supported in part by NIH grants CA51578, CA132831, CA156933, GM069847, and GM095606 (to J.K.). We also acknowledge support of funds in conjunction with grant P30 CA042014 awarded to the Huntsman Cancer Institute, University of Utah. References 1. De Duve, C., De Barsy, T., Poole, B.,

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2 Dendritic Polymers as Targeting Nanoscale Drug Delivery Systems for Cancer Therapy Kui Luo and Zhongwei Gu

2.1 Introduction

Despite the tremendous efforts to deal with cancer, we are still badly in need of a reliable cure for malignant growths. Anticancer chemotherapy drugs, as one of the important manners of cancer therapy, are effective for cancer treatment and tumor suppression. However, there are still some challenges to chemotherapy, such as the nonspecific cytotoxicity, poor water solubility, nonspecific cellular uptake, unfavorable pharmacokinetics, and drug resistance of cancers, which lead to low therapeutic index and high side effects and mortality. To overcome the current shortcomings of chemotherapy, a variety of drug delivery approaches, including polymeric conjugates, polymeric microcapsules, liposomes, and nanoparticles have been extensively investigated as anticancer therapeutics, and they have been approved by either FDA or are in clinical development as cancer treatments. Nanoparticles, as nanoscale drug delivery vehicles [1, 2], are emerging as a class of therapeutics for cancer, suggesting enhanced efficacy, while simultaneously reducing side effects, owing to properties such as higher accumulation in tumors via the enhanced permeability and retention (EPR) effect and active cellular uptake [1, 3]. Among the emergent nanoparticles, liposomes and polymeric nanoscale vehicles have shown great potential clinical impact for the foreseeable future [4]. Currently, natural and synthetic polymeric systems, such as dendritic polymers, have been widely proved to be targeting nanoscale drug delivery vehicles for cancer therapy. Targeting nanoscale drug delivery systems can bring a possible solution to all current problems of low molecular weight drugs, showing promise to expand the therapeutic windows of drugs by increasing delivery to the target tissue as well as the target–nontarget tissue ratio. This will in turn lead to a reduction in the minimum effective dose of the drug. Dendritic polymers are hyperbranched and globular macromolecules with a multivalent surface (nanoscaffold), interior shells, and a core to which the dendrons are attached. The synthetic methodology to construct dendrimers was introduced by Vögtle and his coworkers in 1978 [5]. However, the first series of dendrimers prepared using a well-established method appeared in 1985 with Bioinspired and Biomimetic Polymer Systems for Drug and Gene Delivery, First Edition. Edited by Zhongwei Gu. © 2015 by Chemical Industry Press. All rights reserved. Published 2015 by Wiley-VCH Verlag GmbH & Co. KGaA.

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the introduction of poly(amidoamine) (PAMAM) dendrimers by Tomalia et al. [6]. In addition to PAMAM and poly(propyleneimine) (PPI) dendrimers, peptide dendritic polymers with attractive characteristics are currently under investigation as biomaterials because of their properties similar to proteins, such as good biocompatibility, water solubility, and resistance to proteolytic digestion [7]. Especially during the past few decades, medical science has witnessed the exploration of dendrimers as drug delivery systems. Dendrimers have been proved as a potential alternative route to the development of nanoscale drug delivery vehicles because of their versatility, precise nanostructure, low polydispersity, controllable molecular size, highly adaptable surface chemistry, and chemically stable molecular entities with great flexibility [8–10], which have demonstrated a number of possible advantages, such as drug targeting and pharmacokinetic advantages of typical colloidal or macromolecular delivery systems [10, 11]. The dendritic polymers, such as dendrimers with unique structures, can be employed to either encapsulate or directly conjugate large payloads of anticancer drugs that will be shuttled to the cytoplasm of cancer cells [12]. Compared to linear polymeric analogous, the highly branched, globular architecture of these dendritic macromolecules as nanoscale drug delivery vehicles gives rise to a number of interesting properties, such as increased solubility, very low intrinsic viscosities, and nanoscale size. In addition, dendritic polymer based drug delivery systems have demonstrated significantly higher cellular uptake than that of linear polymeric carriers, such as polyethylene glycol (PEG) [13] and N-(2-hydroxypropyl)methacrylamide (HPMA) [14, 15]. And, the dendritic polymer based drug delivery systems have shown longer blood circulation time, which is benefitting for increasing the drug accumulation in target site (tumor). Those attractive results can be attributed to the unique structural characters, nanoscale size, and compact spherical geometry in solution [16]. In this chapter, we concentrate on dendritic polymers as targeting nanoscale drug delivery systems based on the EPR effect (passive targeting) and on the use of targeting moieties (active targeting) and have been used as delivery drug to tumors for cancer therapy, which are the most numerous and advanced.

2.2 Functional Dendritic Polymers Based Drug Delivery Vehicles for Targeting Tumor Therapy via EPR Effect

It is a well-accepted phenomenon that the endothelial lining of the blood vessel wall in many tumors are more permeable than that in the normal state [17]. In addition, rapid and defective angiogenesis formed from new blood vessels from existing ones leads to a high vascular density in solid tumors [18]. On the basis of those results, in such areas, large gaps exist between endothelial cells in tumor blood vessels and tumor tissues, which show selective extravasation and retention of nanoscale particles. After being administrated via intravenous injection, macromolecules and nanoparticles, such as dendritic polymer based vehicles

2.2

Functional Dendritic Polymers Based Drug Delivery Vehicles for Targeting Tumor Therapy

ranging from 5 to 200 nm in size, can leave the vascular gaps and accumulate inside the interstitial space, which allows 10–30 times higher drug concentration in tumors compared to the blood [19]. The nanoscale drug delivery systems therefore can deliver anticancer drug into the tumor with the increased vascular permeability via EPR effects and allow them to release drugs in the vicinity of the tumor tissue/cells. Because the cut-off size of the permeabilized vasculature varies from case to case, the spontaneous “passive” targeting efficacy can be optimized via the control size of dendritic polymer nanoscale vehicles. It is now a well-established phenomenon that the diameters of nanoscale vehicles should be less than 200 nm to avoid uptake by the reticuloendothelial system (RES), which is effective for passive targeted drug delivery to solid tumors [20]. In addition to the cut-off size of nanoparticles, for successful EPR-mediated-targeting, this type of targeting requires dendritic polymer based drug delivery systems to have long circulation in the blood to extended periods of time, resulting in a sufficient level of accumulation in the tumor tissue. The circulation time of drug delivery system in the blood should be more than 6 h [21, 22]. However, for most of the dendritic polymers, such as PAMAM and PPI dendrimers, the size is below 10 nm and the particles showed positive charge [23]. That’s attributed to their rich amino groups in their structures (primary amine at the surface and tertiary amine groups in the branch). To address those shortcomings, the dendritic polymers were usually modified with certain negative charged moieties and water soluble polymers with a well-solvated and flexible main chain, such as PEG [24]. The surface-grafted ligands rapidly increased the size of dendritic polymer. Simultaneously, the “protective” moieties effectively prevent the protein adsorption of drug carriers and their clearance by the RES, demonstrating longer blood circulation and enhanced accumulation into tumor [25, 26]. Although passive targeting approach is one of the basis of cancer therapy, it is important to mention that it is still not very effective for cancer therapy because only a small portion of the administrated drug can accumulate in the solid tumor tissues. In different tumor situations the integrity of vascular endothelium remains unaffected and there is no opportunity for EPR effects, as the size of gaps between tumor blood vessels and tumor tissues is different at different phases of the cancer. Therefore, passive targeting approach for cancer therapy so far is not universal. It is known that solid tumors and tumor angiogenesis often overexpress specific antigens, receptors on cell surfaces or integrins, such as αv β3 integrins [27–31], which serve as the important target sites of prepared drug delivery vehicles. Until now, the most natural and universal way to enhance the nonspecific drug affinity toward tumor tissues/cells is directly introducing the drug and a tumor targeting moiety to a single vehicle, giving a targeting drug delivery system. The targeting moiety with the capability of specific recognition and binding to a target sites (biomarkers) enables the functionalized drug delivery vehicle for active tumor targeting. The following substances – peptide, mono-, oligo-, and polysaccharides, antibodies and their fragments, folate, and so on – can be used as targeting moieties [32]. In general, direct coupling of drugs and targeting moieties to a vehicle seems the simplest way to prepare a targeting drug delivery system. However, one

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2 Dendritic Polymers as Targeting Nanoscale Drug Delivery Systems for Cancer Therapy

of the challenges is to control the ratio of pharmaceutical agents and targeting moieties conjugated onto a single vehicle. The interactions of isolated targeting moiety-receptor are typically very weak. One solution for the weakness of these isolated interactions can be achieved through the synthetic cluster targeting moieties on a dendritic framework interacting with their receptors, resulting in a linear increase of affinity for multimeric receptors on cancer cell surfaces [33]. That is attributed to the multivalency of dendrimer [34]. The significantly increased targeting capability manifested by dendrimer has been termed as “cluster effect” [35], which is the feature of dendritic polymer used as targeting drug delivery for cancer therapy. Previous studies, for example, have demonstrated that glycopeptide dendrimers bearing multiple galactose moieties on a single dendritic scaffold showed excellent inhibitory capacities for asialoglycoprotein receptors (ASGP-Rs) [36]. 2.2.1 Functional Dendritic Polymers for Encapsulation of Anticancer Drugs

Active anticancer pharmaceutical agents can be loaded into the internal voids of dendritic polymers and assembled with polymers by physical incorporation, or by covalent conjugation onto the dendritic architecture [37]. The functional vehicles, especially PEGylation dendritic polymers, for drug loading by physical incorporation showed great possibility as drug delivery for cancer therapy. The unique chemical architecture of dendrimers with a well-defined core, backbone, and multivalent periphery has been firstly promoted as a candidate for host–guest chemistry for drug delivery. The dendritic structures and anticancer drugs are assembled into nanoscale particles. Initial studies of dendrimers as potential delivery anticancer agents focused on their use for noncovalent encapsulation (Figure 2.1a), which is to enhance drug solubility and bioavailability and act as release modifiers and platforms for drug targeting. The drug encapsulation mediated by dendritic polymers is based on physical entrapment and interactions between drug ionic groups and oppositely charged dendritic surfaces. For dendrimers, several types of interactions have been explored, which can be broadly subdivided into the entrapment of drugs within the dendritic architecture (involving nonbonding interactions, such as electrostatic, hydrophobic, and hydrogen bond interactions) and the interaction between a drug and the surface of a dendrimer (electrostatic interactions). The drug loading capacity is influenced by several factors, such as nanoenvironments, generation, and molecular weight of dendrimer, characteristics of architecture and internal cavities, molecular weight, and structure of drug. In 1982, Maciejewski had shown the use of egg shell-like architectures for the encapsulation of guest molecules in polymers [38]. Until now, several types of dendritic polymers, including PAMAM and PPI dendrimers, have been investigated for the encapsulation of drugs via hydrogen bonding or electrostatic interactions and have been designed for triggered release. The dendrimer with high generation showed higher drug encapsulation capability and could assemble into nanoparticles as passive targeting. However, it’s not simple to prepare high generation dendrimer because of the high steric hindrance

2.2

Functional Dendritic Polymers Based Drug Delivery Vehicles for Targeting Tumor Therapy

(a)

(b) : Drugs

: Linker

Figure 2.1 The illustration of dendrimer-based drug delivery vehicles carrying drug via noncovalent encapsulation (a) and covalent conjugation (b).

to chemical reaction accompanied with the increasing of generation, especially for the high generation one (beyond 5). The emergence of click chemistry has provided a powerful tool to address the challenges because of its high efficiency, quantitative yields, and technical simplicity, which have been utilized for the preparation of various dendrimers with high generations. Simultaneously, the heterocyclic residues formed from click reaction of azide–alkyne cycloaddition are structurally adjacent to amides (H-bond acceptors or H-bond donors), can bind with drugs in a remarkably specific manner [34], which was used to improve the dendrimer-drug binding stability by providing groups that can act as H-bond acceptors and increase hydrophobic interactions of these carriers with drugs in aqueous solution. However, for the high generation PAMAM and some dendrimers synthesized via click chemistry, in vitro studies have shown that those cationic dendrimers may be cytotoxic because of their excess protonated groups and hemolytic effects, resulting in high nonspecific uptake by the RES particularly in the liver and lungs, reducing the relative drug delivery system accumulation in tumor tissue. In order to overcome the shortcomings, one possible strategy is to functionalize the surface of the dendrimer. The dendrimers with high charge density can be altered via the tunable surface functional groups, which can optimize the local environment and influence cellular interactions and cell uptake. The PAMAM dendrimers were functionalized with amine, hydroxyl, and carboxyl groups, showing different endocytosis uptake mechanisms. Anionic dendrimers functionalized with carboxyl groups are partly taken up by caveolae, while the cationic and neutral dendrimers appear to be taken in by a nonclathrin, noncaveolae mediated mechanism because of the electrostatic interactions and other nonspecific fluid-phase endocytosis [39]. Therefore, it is possible to modulate the cell entry kinetics and target therapeutic agents to specific cell organelles by engineering the surface charge on the dendrimer. Comparing the biodistribution of cationic PAMAM dendrimers and their neutral counterparts synthesized by partial or full acetylation of the surface amine groups in nude mice bearing melanoma and prostate tumors, both

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2 Dendritic Polymers as Targeting Nanoscale Drug Delivery Systems for Cancer Therapy

dendrimers resulted in a similar distribution profile to all major organs within 1 h post-injection with particularly high accumulation in the lungs, kidneys, and liver (27.9–28.6% ID/g). Despite similar biodistribution profiles manifested by both cationic and neutral dendrimers, cationic dendrimers showed higher net accumulation in each organ. That is attributed to their favorable electrostatic interaction between dendrimer and the negatively charged epithelial and endothelial cell surface. It is notable that anionic PAMAM-COOH dendrimers showed high distribution to the liver and quick elimination into the urine. However, less than 5% of the initial dose remained in the systemic circulation 24 h after administration. That is because of the dendrimer’s small hydrodynamic size. The short blood circulation time reduced the drug delivery efficacy to target site (tumor) and anticancer efficacy, despite possible passive targeting. In addition, it is reported that drugs loaded noncovalently into dendrimers, even the dendrimer with high generation, are released rapidly, rendering them ineffective for targeted drug delivery. It is generally accepted that the passive targeting efficiency, pharmacokinetics, and biodistribution also depend on the size of the nanoscale drug delivery systems. Previous studies have shown that particles smaller than 5.5 nm primarily undergo renal clearance, whereas particles of intermediate size are more variable and those >12 nm primarily undergo hepatic clearance. However, the size of reported dendrimers is 5–10 nm, which is not satisfactory as drug delivery vehicle for cancer therapy. To improve the encapsulation efficacy, small size and release properties remained as the major challenges in dendrimer-based delivery systems, the dendrimer or dendron were functionalized with polymer, giving functional dendritic polymers, which have shown significant influence on their performance as drug delivery vehicles and in particular on encapsulation efficiency and release. For example, polyether-co-polyester (PEPE) dendrimer with polyethylene oxide (PEO) in the interior cavity was prepared not only to increase the size of the cavity of dendrimer but also to enhance dendrimer’s ability to encapsulate drug by virtue of its solublizing properties. Hildgen et al. also reported a series of PEPE dendrimers of different architecture [40]. The studies on biocompatibility of the synthesized dendrimers on RAW 264.7 cells using lactate dehydrogenase (LDH) assay showed no cytotoxicity even at the concentration of 250 mg/ml, suggesting that they are acceptable as drug delivery vehicle. The functional dendrimers showed good capacity to encapsulate methotrexate (MTX) because of the high drug loading (24.5% w/w). The encapsulation capacity can be improved by increasing the number of branches and the size of internal voids, whereas the mechanisms of encapsulation may be because of the physical entrapment, weak hydrogen bonding, and hydrophobic interactions. In addition, for this kind of dendrimers, the release of MTX was biphasic because of the burst release in 6 h followed by a slower release over a period of 50 or 168 h. It is notable that the burst drug release can be overcome by increasing the number of the branches of dendrimer. Therefore, the drug encapsulation capability and drug release characters can be controlled by inputs into the design and optimization of chemical structure of dendritic nanocarriers. PPI dendrimers were functionalized with

2.2

Functional Dendritic Polymers Based Drug Delivery Vehicles for Targeting Tumor Therapy

polysorbate 80 (P80) to encapsulate anticancer drug docetaxel (DTX), giving nanoscale drug delivery systems (DTX-P80-PPI) for targeting the brain tumor [41]. In vitro cytotoxicity studies of free DTX, DTX loaded PPI dendrimer (DTXPPI), and DTX-P80-PPI dendrimers on U87MG human glioblastoma cell line showed the dendrimers based system have not resulted in any significant cytotoxic effect. However, the in vivo anticancer activity in brain tumor bearing rats revealed that DTX loaded dendrimers functionalized with P80 (DTX-P80-PPI) reduced the tumor volume extremely significantly as high as 50%. The median survival time for brain tumor bearing mice treated with DTX-P80-PPI dendrimers was 42 days, which was extended more significantly compared to DTX-PPI (23 days) and free DTX (18 day). The targeting efficiency and higher biodistribution of polysorbate functionalized dendrimer into the brain were confirmed via gamma scintigraphy and biodistribution studies. Those results concluded that the developed dendrimer functionalized with polymers has potential to deliver significantly higher amount of drug to tumor for improving therapeutic indexes. That is attributed to the increased size and higher drug encapsulation, which enhances the passive targeting efficiency and drug accumulation in tumor. It should be noted that PEG has been used to modify the dendrimer with small size to enhance the solubilization and controlled release of chemotherapeutics for cancer therapy. To enhance the blood circulation of dendrimers, PEG, which can extend retention time in blood by decreasing nonspecific interactions with endogenous components and macrophages, has been conjugated to dendrimers to form a new kind of dendritic structures. The PEGylated dendrimers have demonstrated better retention in blood, low accumulation in organs, and high accumulation in tumor tissue via the EPR effect [32]. On the basis of the above observation, Margerum et al. reported PEGylated PAMAM dendrimers as magnetic resonance imaging (MRI) contrast agents. The macrocycle 1-(4-isothiocyanatobenzyl)amido-4,7,10-triacetic acid-tetraazacyclododecane (D03A-bz-NCS) was conjugated to the terminal amine sites of starburst PAMAM dendrimers (G2 and G3), giving dendritic polychelates. Gadolinium ion (Gd3+ ) was chelated to the polychelates, creating water soluble, monodisperse, and dendritic MRI contrast agents. Postinjection into rat, there was an increase in blood elimination half-life with molecular weight ranging from 11 min for lower molecule agent (MW: 22 kDa) to 115 min for the higher molecule ones (MW: 61.8 kDa). We also reported the mPEGylated peptide dendrimer as MRI contrast agents, as shown in Figure 2.2 [42]. For mPEGylated dendrimer, there was a higher Gd3+ concentration over time than the lower molecular weight dendrimers in the blood. Intravenous administration of un-mPEGylated dendrimers (G3) were rapidly eliminated from the circulation within a few minutes. In contrast, the concentration of PEGylated dendrimer was about 38 μg Gd3+ /g tissue at 1 h postinjection, 30 times higher than other un-PEGylated formulations. These results indicate that the Gd3+ -labeled peptide dendrimer with methoxy polyethylene glycol (mPEG) modification would be potentially used as efficient MRI probes because of its prolonged retention in the circulation and decreased accumulation in the organs. It

35

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2 Dendritic Polymers as Targeting Nanoscale Drug Delivery Systems for Cancer Therapy

G3-12Gd-DTPA-10mPEG

G3-24Gd-DTPA

G3-12Gd-DTPA

G3-6Gd-DTPA

G3-24Gd-DTPA

μg Gd/g tissue

80

G3-12Gd-DTPA

40 20

G3-12Gd-DTPA-10mPEG

G3-6Gd-DTPA

0

O O

= mPEG2000

(a)

60

= −COOH

N O

O O

N

O

HOH

10

N

Gd O

O O

O

(b)

Figure 2.2 The structures of gadoliniumbased peptide dendrimers (a) and the PEGylated dendrimers showed long blood circulation after administration of peptide

20

30

60

Time (min) dendrimers at a dose of 0.05 mmol Gd3+ /kg in BALB/C mice (b). Adapted from Ref. [42], reproduced by permission of the Elsevier Publishing Group.

is worth noticing that the dendrimer generation and PEG molecular weight influence the drug encapsulation efficacy, as Kojima et al studies on the relationship between dendrimer generation (G3 and G4) and PEG molecular weight (550 or 2000) and the ability of PEG grafted dendrimers to encapsulate the hydrophobic drugs adriamycin (II) and MTX (III) [43]. Those results suggest that PEGylation will be an effective strategy for improving the in vivo performance of dendrimers. In addition, because PEG is highly water soluble and biocompatible, it was typically conjugated to the surface of a dendrimer to provide a hydrophilic shell around a hydrophobic dendritic core to form a unimolecular micelle, which drastically augments drug loading and solubilization. PEGylated dendrimers also eliminate the naked dendrimeric scaffold drawbacks of hemolytic toxicity (especially to cationic ones), uncontrolled drug burst release, macrophageal uptake, short blood half-life, and so on. For instance, the surface of the six generation dendrimers based on melamine was modified with amine, guanidine, carboxylate, sulfonate, phosphonate, and PEG, respectively, and the hemolytic potential and cytotoxicity were evaluated. The results demonstrated the cationic dendrimers were more cytotoxic and hemolytic than anionic or PEGylated dendrimers. Also, the PEGylated dendrimer showed no toxicity, lethality, or abnormalities in blood chemistry [44]. PEGylated dendrimers can be used as liposomes, which have been shown to sustain drug release. Gardikis et al. reported the PEGylated G2 dendrimer showed that at 96 h cumulative doxorubicin (DOX) release was 27.9 ± 2.8%. In contrast, 74.6 ± 7.8% drug was released [45]. Bhadra et al. reported PEGylated PAMAM dendrimers for the delivery of 5-fluorouracil (5-FU) [46]. The results showed this dendritic system functionalized with PEG (5 kDa) may be a suitable vehicle for the extended delivery of anticancer drugs in vitro as well as in vivo. This dendritic system demonstrated 12-fold higher loading capacity in PEGylated G4 PAMAM dendrimers and 6-fold lower release rate compared to

2.2

Functional Dendritic Polymers Based Drug Delivery Vehicles for Targeting Tumor Therapy

non-PEGylated dendrimers. The blood level of the system was prolonged, as it was detectable for up to 12 h. The average release rate was found to be 0.679%, which was nearly one-sixth of that from non-PEGylated systems. Importantly, the dendritic system showed low hemolytic toxicity. The hemolysis of red blood cells (RBCs) declined significantly may be because of the inhibition of the interaction of RBCs with the charged quaternary ammonium ion by PEG chains, because the surface and the PAMAM dendrimer were protected. Therefore, the use of PEGylated dendrimers in cancer treatment technology has been studied extensively. 2.2.2 Chemical Conjugation Functional Dendritic Polymers as Drug Delivery Systems

Physical incorporation of drugs using dendritic polymers has some straightforward advantages, rapid, and simple preparation without adversely affecting drug pharmacological activity. The encapsulation of drug in the dendritic voids protects labile molecules from degradation and also eliminates solubility issue associated with hydrophobic compounds. Despite these potential advantages, in general, the encapsulation approach has met with problems, including low stability in terms of storage and burst release of many chemotherapeutics in plasma, variation of the concentration of the solubilized drug from batch to batch, and low drug loading capacity, often leading to less pharmacokinetic or therapeutic benefit compared with the administration of free drug. In addition, in vitro release of physically encapsulated drugs from dendritic cavity is usually rapid, which is manifested by several factors, such as the drug partition coefficient between hydrophobic and aqueous environments, strength of drug/dendritic polymer interactions, molecular weight of dendritic polymer, and modified surface groups. To overcome the limitations caused by encapsulation, chemical conjugation serves to prepare dendritic polymer based drug delivery systems (Figure 2.1b). Anticancer drugs were covalently conjugated to dendritic scaffold to achieve increased drug loading and controlled spatial and temporal release of the attached drugs. Using a chemical conjugation approach will also bring about selective drug release and enhanced biosafety in vivo, as shown in Figure 2.3 [47]. For instance, three different DOX formulations (DOX in saline, DOX-polylysine dendrimer conjugates, a stealth liposome encapsulated DOX) were compared and studied. The pharmacokinetics, biodistribution, and antitumor efficacy of three DOX preparations were investigated in Walker 256 tumor bearing rats. Liposomal and dendritic conjugate delivery systems demonstrated more prolonged plasma exposure of total DOX when compared to the administration of free drug DOX in saline. Biodistribution profiles revealed enhanced accumulation of the two drug delivery systems in tumors when compared to free drug DOX alone. All three DOX formulations resulted in similar in vivo anticancer efficacy. However, compared to dendritic conjugates, markers of in vivo systemic toxicity (spleen weight, white blood cell counts, body weight, and cardiotoxicity) were observed in rats administrated with DOX and liposomal DOX. The preliminary results suggest that dendritic conjugates display

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2 Dendritic Polymers as Targeting Nanoscale Drug Delivery Systems for Cancer Therapy

Doxorubicin (DOX) solution

DOX Systemic toxicity

Dox-dendrimer Dox-liposome

Control

Tumour size

38

Dendrimer-DOX Doxorubicin Formulations 0

5

10

15

20

Reduced systemic toxicity

25

Equivalent tumouricidal activity Figure 2.3 Compared to doxorubicin, the dendrimer-doxorubicin conjugate demonstrated much higher anticancer efficacy as well as a reduced incidence of cardiotoxicity. The results also suggest that dendrimer-based

vehicle may provide advantage over PEGylated liposomal formulation of doxorubicin via a reduction in systemic toxicity. Adapted from Ref. [47], reproduced by permission of the American Chemical Society.

similar antitumor efficacy to PEGylated liposomal DOX, but with lower systemic toxicity (resulting from reduced drug exposure to nontarget organs) [48]. If the drug was conjugated to the dendritic scaffold via stimuli-responsive bonds, drug controlled release can be achieved by the changes in the biological microenvironment, such as variation in pH, temperature, or concentration of a specific enzyme. The large number of surface groups and the versatility in their chemical structures enable the possibility of conjugating different anticancer drugs, imaging agents, tumor targeting ligands while maintaining the compact spherical nanoscale particles in solution, resulting in targeting drug delivery systems. In the early 1990s, Barth and coworkers reported that boronated monoclonal antibodies were conjugated to a dendrimer via stable urea linkages and utilized those conjugated dendritic systems for neutron capture therapy. The preliminary anticancer efficacy was observed because of the caused necrosis of neighboring cancer cells via localized neutron ionization [49, 50]. Duncan and coworkers also coupled cisplatin (Pt), a hydrophobic DNA intercalating agent, to PAMAM dendrimer via an ester linkage, resulting in Pt-loaded dendritic systems where the Pt content was 20–25 wt% [51]. Despite the high aqueous solubility and stability (10-fold higher aqueous solubility compared to free Pt), this drug delivery system did not produce the desired anticancer activity. Specifically, the system displayed insignificant toxicity toward three cancer cells lines when treated with 0.1 × 10−5 to 0.01 mg/ml Pt equivalent for 72 h. That is attributed to the too slow drug release as this system on incubation in phosphate buffer saline (PBS) (pH 7.4) and citrate buffers (pH 5.5) at 37 ∘ C for 72 h displayed great stability (200 nm [19, 20], in which the micelles with optimum size range of less than 100 nm are considered to favor tumor accumulation of drugs via EPR effect, on which the micellar formulations mainly rely including NK012, NK105, and NC6004 under clinical trials. Some nanoparticle formulations such as Doxil with diameters larger than 100 nm have exhibited limited penetration and accumulation in hypovascular and hypopermeable tumors, and thus particle size smaller than 100 nm potentially plays a key role in tumor penetration capacity of the micelles. Recently, Cabral et al. [21] explored the tumor accumulation and thereof therapeutic efficacy of polymeric micelles with various diameters (e.g., 30, 50, 70, and 100 nm) against both highly and poorly permeable tumors (Figure 4.3). 1,2-diaminocyclohexaneplatinum(II) (DACHPt)-loaded micelles

89

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4 Polymeric Micelles for Cancer-Targeted Drug Delivery

O H N

O nN

H

m

H

H2 N

COO Pt COO

+−

Na O

O

N H2

CHOO Pt H2N

H2 N

H2O

NH2

COO− COO−

Cl Pt

Cl

N H2

Poly(ethylene glycol)-b-poly(glutamic acid) (PEG-b-PGlu) copolymer + O H N N H

H

m

Na+ −O

Self-assembly in aqueous environment

In chloride ioncontaining media

O

Poly(glutamic acid) (PGlu) homopolymer + H 2

N

H2O Pt2+

N H2

(a)

H2O

(1,2-Diaminocyclohexane)platinum(II) (DACHPt) aqueous complex No micelle

30 nm

50 nm

70 nm

100 nm

C26

BXPC3

Nest

(b)

Fibrotic tissue

Figure 4.3 (a) Schematic nanostructure of DACHPt-loaded micelles. In aqueous media containing chloride ions, DACHPt (yellow circles) can be released from micelles through ligand exchange. (b) Histological examination of C26 tumor and BxPC3 tumor by H & E staining and fluorescent microscopic images of C26 and BxPC3 tumor sections at

24 h post-injection after intravenous administration of fluorescent micelles with various particle size. Micelles were labeled with Alexa 594 (red). Blood vessels were marked with PECAM-1 and Alexa 488 secondary antibody (green) [21]. (Reproduced with permission from Ref. [21]. )

4.3

Particle Size of Micelles

are constructed using PEG-b-PGlu (PGlu: poly(glutamic acid)) copolymer and PGlu homopolymer. All the DACHPt-loaded micelles with different particle size can penetrate through highly permeable tumors in mice, but only 30 nm micelles can penetrate into poorly permeable BxPC3 pancreatic tumors and exhibit a superior antitumor efficacy. It indicates that the micelles exhibit no size-dependent restrictions on extravasation and penetration in hypervascular tumors. However, only micelles with size smaller than 50 nm can penetrate poorly permeable hypovascular tumors. In addition, it offers a new strategy to improve anticancer efficacy of the micelles by enhancing their extravasation and penetration through the neovasculature using a transforming growth factor-β inhibitor (TGF-β-I). Schadlich et al. [22] also found the size-dependent in vivo behavior of the micelles. The well-defined micelles (111 nm in diameter) accumulate more efficiently at tumor when compared with the micelles with the average diameter of 166 nm. Particle size plays a key parameter for optimizing in vivo accumulations of anticancer drugs at tumors. So far, there have been only a few studies about the relationships between particle size and anticancer efficacy. It is highly necessary to develop a precise assembly of micelles with tunable size and explore the relationships on different tumor models with various pathological conditions, which may facilitate the optimization of micellar formulations for achieving superior anticancer efficacy [22]. On the other hand, particle size is also a key parameter affecting the endocytosis pathways of nanoparticles, which can generally have obvious influence on their intracellular delivery. Large particles (>1 μm) can be internalized by phagocytosis, which are generally present on phagocytic cells such as macrophages, neutrophils, and dendritic cells. But the pinocytosis is more relevant to cell internalization of nanoparticles and have several pathways including clathrin-, caveolae-, RhoA-, CDC42-, ARF6-, Flotillin-mediated endocytosis, and macropinocytosis, which display various size limitations and can likewise lead to various intracellular trafficking (Figure 4.4). For instance, clathrin-mediated endocytosis will encounter the lysosomal degradation of the cargo, whereas caveolae-mediated endocytosis has no digestive environment [23]. Therefore, particle size may potentially regulate intracellular trafficking of the micelles and result in diverse response of anticancer efficiency. Generally, the micelles are reported to undergo clathrin- or caveolae-mediated endocytosis. However, it is difficult to estimate endocytosis pathway of micelles according to particle size limitation. For example, Pluronic P85 micelles with particle size of 14.6 nm exhibit a clathrin-endocytosis, whereas poly(methyl methacrylate)-b-poly(polyethylene glycol methyl ether methacrylate) micelles with an average size of 20 nm undergo caveolae-mediated endocytosis [24]. In some case, the micelles can undergo a major pathway, accompanying with several other pathways as well [25]. To date, there have been only a few studies about the influence of particle size on the endocytosis pathways of micelles. The precise control of micellar size in the range of 10–100 nm remains a major challenge for elucidating the relationship between size and endocytosis pathways [22].

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4 Polymeric Micelles for Cancer-Targeted Drug Delivery Phagocytosis (Particle-dependent)

Macropinocytosis (