Biomedical applications of electroactive polymer actuators

Biomedical Applications of Electroactive Polymer Actuators

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

Biomedical Applications of Electroactive Polymer Actuators FEDERICO CARPI University of Pisa, Pisa, Italy ELISABETH SMELA University of Maryland, College Park, USA

A John Wiley and Sons, Ltd., Publication

This edition first published 2009 Ó 2009 John Wiley & Sons Ltd. Registered office John Wiley & Sons Ltd, The Atrium, Southern Gate, Chichester, West Sussex, PO19 8SQ, United Kingdom. For details of our global editorial offices, for customer services and for information about how to apply for permission to reuse the copyright material in this book please see our website at www.wiley.com. The right of the author to be identified as the author of this work has been asserted in accordance with the Copyright, Designs and Patents Act 1988. All rights reserved. No part of this publication may be reproduced, stored in a retrieval system, or transmitted, in any form or by any means, electronic, mechanical, photocopying, recording or otherwise, except as permitted by the UK Copyright, Designs and Patents Act 1988, without the prior permission of the publisher. Wiley also publishes its books in a variety of electronic formats. Some content that appears in print may not be available in electronic books. Designations used by companies to distinguish their products are often claimed as trademarks. All brand names and product names used in this book are trade names, service marks, trademarks or registered trademarks of their respective owners. The publisher is not associated with any product or vendor mentioned in this book. This publication is designed to provide accurate and authoritative information in regard to the subject matter covered. It is sold on the understanding that the publisher is not engaged in rendering professional services. If professional advice or other expert assistance is required, the services of a competent professional should be sought. The publisher and the author make no representations or warranties with respect to the accuracy or completeness of the contents of this work and specifically disclaim all warranties, including without limitation any implied warranties of fitness for a particular purpose. This work is sold with the understanding that the publisher is not engaged in rendering professional services. The advice and strategies contained herein may not be suitable for every situation. In view of ongoing research, equipment modifications, changes in governmental regulations, and the constant flow of information relating to the use of experimental reagents, equipment, and devices, the reader is urged to review and evaluate the information provided in the package insert or instructions for each chemical, piece of equipment, reagent, or device for, among other things, any changes in the instructions or indication of usage and for added warnings and precautions. The fact that an organization or Website is referred to in this work as a citation and/or a potential source of further information does not mean that the author or the publisher endorses the information the organization or Website may provide or recommendations it may make. Further, readers should be aware that Internet Websites listed in this work may have changed or disappeared between when this work was written and when it is read. No warranty may be created or extended by any promotional statements for this work. Neither the publisher nor the author shall be liable for any damages arising herefrom. Library of Congress Cataloging-in-Publication Data Biomedical applications of electroactive polymer actuators / [edited by] Federico Carpi, Elisabeth Smela. p. ; cm. Includes bibliographical references and index. ISBN 978-0-470-77305-5 (H/B) 1. Polymers in medicine. 2. Conducting polymers 3. Actuators. I. Carpi, Federico, 1975– II. Smela, Elisabeth. [DNLM: 1. Polymers—diagnostic use. 2. Polymers—therapeutic use. 3. Biomedical Technology. 4. Equipment and Supplies. QT 37.5.P7 B6147 2009] R857.P6B5485 2009 610.280 4—dc22 2008056029 A catalogue record for this book is available from the British Library. ISBN: 978-0-470-77305-5 (H/B)

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Contents

Preface List of Contributors Introduction SECTION I 1

page xv xvii 1

POLYMER GELS

5

Polymer Gel Actuators: Fundamentals Paul Calvert

7

1.1 1.2

1.3 1.4

1.5 1.6

1.7

Introduction and Historical Overview Properties of Gels 1.2.1 Biological Gels 1.2.2 Mechanical Properties of Simple, Single-Phase Gels 1.2.3 Elastic Moduli 1.2.4 Strength 1.2.5 Multi-Phase Gels 1.2.6 Double Network Gels 1.2.7 Transport Properties 1.2.8 Drying Chemical and Physical Formation of Gels Actuation Methods 1.4.1 Thermally Driven Gel Actuators 1.4.2 Chemically Driven Gel Actuators 1.4.3 Gels Driven by Oscillating Reactions 1.4.4 Light Actuated Gels 1.4.5 Electrically Driven Gel Actuators 1.4.6 Electro- and Magneto-Rheological Composites 1.4.7 LC Elastomers Performance of Gels as Actuators Applications of Electroactive Gels 1.6.1 Gel Valves and Pumps 1.6.2 Light Modulators 1.6.3 Gel Drug Delivery 1.6.4 Gel Sensors Conclusions References

7 8 8 9 10 10 12 13 14 15 16 19 19 20 22 23 23 25 26 26 30 30 30 31 32 32 33

vi

2

Contents

Bio-Responsive Hydrogels for Biomedical Applications Tom McDonald, Alison Patrick, Richard Williams, Brian G. Cousins and Rein V. Ulijn

43

2.1 2.2 2.3 2.4 2.5

Introduction Chemical Hydrogels Physical Hydrogels Defining Bio-Responsive Hydrogels Bio-Responsive Chemical Hydrogels 2.5.1 Actuation Based on Changing the Cross-Linking Density 2.5.2 Actuation Based on Changes in Electrostatic Interactions 2.5.3 Actuation Based on Conformational Changes Bio-Responsive Physical Hydrogels 2.6.1 Enzyme-Responsive Physical Hydrogels Electroactive Chemical Hydrogels Conclusion References

43 44 44 44 46

Stimuli-Responsive and ‘Active’ Polymers in Drug Delivery Aram Omer Saeed, Jo´hannes Pa´ll Magnu´sson, Beverley Twaites and Cameron Alexander

61

3.1 3.2

61 62 62 63 63

2.6 2.7 2.8

3

3.3

3.4

Introduction Drug Delivery: Examples, Challenges and Opportunities for Polymers 3.2.1 Oral Drug Delivery Systems 3.2.2 Parenteral Drug Delivery 3.2.3 Topical and Transdermal Drug Delivery 3.2.4 Delivery Challenges for Biomolecular Drugs and Cell Therapeutics 3.2.5 Peptides and Proteins 3.2.6 Nucleic Acids 3.2.7 Cell Delivery Emerging State-of-the-Art Mechanisms in Polymer Controlled Release Systems 3.3.1 Technologies for Controlled Drug Release 3.3.2 Polymer–Drug Conjugates 3.3.3 Polymer–Protein Conjugates 3.3.4 Polymer–Nucleic Acid Conjugates 3.3.5 Polymer–Nucleic Acid Complexes Responsive or ‘Smart’ Polymers in Drug Delivery 3.4.1 Soluble Smart Polymers 3.4.2 Responsive Polymer–Drug Conjugates 3.4.3 Responsive Polymer–Protein Conjugates 3.4.4 Responsive Polymers for DNA Delivery

46 49 51 53 53 56 57 57

64 64 65 65 67 67 67 67 68 68 73 73 76 76 77

Contents

3.5 3.6

4

Recent Highlights of Actuated Polymers for Drug Delivery Applications Conclusions and Future Outlook References

Thermally Driven Hydrogel Actuator for Controllable Flow Rate Pump in Long-Term Drug Delivery Piero Chiarelli and Pietro Ragni 4.1 4.2 4.3

4.4 4.5

Introduction Materials and Methods Hydrogel Actuator 4.3.1 Thermo-Mechanical Gel Dynamics 4.3.2 Experimental Results Pump Functioning Conclusion References

SECTION II 5

78 80 81

89 89 90 90 91 93 97 98 98 101

IPMC Actuators: Fundamentals Kinji Asaka and Keisuke Oguro

103

5.1 5.2

103 104 104 105 108 110 113 116 117 118

5.3 5.4 5.5 5.6 5.7

6

IONIC POLYMER–METAL COMPOSITES (IPMC)

vii

Introduction Fabrication 5.2.1 Ionic Polymer 5.2.2 Plating Methods Measurement Performance of the IPMC Actuator Model Recent Developments Conclusion References

Active Microcatheter and Biomedical Soft Devices Based on IPMC Actuators Kinji Asaka and Keisuke Oguro 6.1 6.2 6.3 6.4

6.5

Introduction Fabrication of the IPMC Device Applications to the Microcatheter Other Applications 6.4.1 Sheet-Type Braille Display 6.4.2 Underwater Microrobot 6.4.3 Linear Actuators for a Biped Walking Robot Conclusions References

121 121 122 124 127 127 130 134 135 135

viii

7

Contents

Implantable Heart-Assist and Compression Devices Employing an Active Network of Electrically-Controllable Ionic Polymer–Metal Nanocomposites Mohsen Shahinpoor 7.1 7.2 7.3 7.4 7.5 7.6 7.7 7.8 7.9 7.10

8

9

Introduction Heart Failure Background of IPMNCs Three-Dimensional Fabrication of IPMNCs Electrically-Induced Robotic Actuation Distributed Nanosensing and Transduction Modeling and Simulation Application of IPMNCs to Heart Compression and Assist in General Manufacturing Thick IPMNC Fingers Conclusions References

IPMC Based Tactile Displays for Pressure and Texture Presentation on a Human Finger Masashi Konyo and Satoshi Tadokoro

137 137 139 140 141 142 144 146 149 155 157 157

161

8.1 8.2 8.3 8.4 8.5 8.6

Introduction IPMC Actuators as a Tactile Stimulator Wearable Tactile Display Selective Stimulation Method for Tactile Synthesis Texture Synthesis Method Display Method for Pressure Sensation 8.6.1 Method 8.6.2 Evaluation 8.7 Display Method for Roughness Sensation 8.7.1 Method 8.7.2 Evaluation 8.8 Display Method for Friction Sensation 8.9 Synthesis of Total Textural Feeling 8.9.1 Method 8.9.2 Experiments 8.10 Conclusions References

161 162 164 165 167 168 168 168 169 169 170 171 172 172 172 173 173

IPMC Assisted Infusion Micropumps Il-Seok Park, Sonia Vohnout, Mark Banister, Sangki Lee, Sang-Mun Kim and Kwang J. Kim

175

9.1 9.2 9.3

175 176 177 178

Introduction Background of IPMCs Miniature Disposable Infusion IPMC Micropumps 9.3.1 Configuration of the IPMC Infusion Pump

Contents

9.4

9.5

9.3.2 The Control System 9.3.3 Performance Testing Modelling for IPMC Micropumps 9.4.1 Equivalent Bimorph Beam Model for IPMC Actuators 9.4.2 IPMC Diaphragm Conclusions References

SECTION III 10

180 181 181 181 182 189 189 193

Conjugated Polymer Actuators: Fundamentals Geoffrey M. Spinks, Gursel Alici, Scott McGovern, Binbin Xi and Gordon G. Wallace

195

10.1 10.2 10.3 10.4 10.5

195 197 200 201

10.6 10.7

10.8

11

CONJUGATED POLYMERS

ix

Introduction Molecular Mechanisms of Actuation in ICPs Comparison of Actuation Performance in Various ICPs Electrochemistry of ICPs Effect of Composition, Geometry and Electrolyte on Actuation of PPy 10.5.1 Effect of the Dopant Ion 10.5.2 Effect of Solvent 10.5.3 Charge Transfer Processes 10.5.4 Effect of Porosity/Morphology Mechanical System Response Device Design and Optimization 10.7.1 How to Tailor Actuator Performance to Meet Design Requirements 10.7.2 Design of a Swimming Device 10.7.3 Device Testing Future Prospects References

204 204 206 208 212 212 217 217 219 221 222 223

Steerable Catheters Tina Shoa, John D. Madden, Nigel R. Munce and Victor X.D. Yang

229

11.1 11.2 11.3

229 229 231 231 232 232 232 234 234 235

11.4

Introduction Catheters: History and Current Applications Catheter Design Challenges 11.3.1 Biocompatibility 11.3.2 Small Size 11.3.3 Low Cost 11.3.4 Structural Rigidity Active Steerable Catheters 11.4.1 Non-EAP Based Steerable Catheters 11.4.2 EAP Based Steerable Catheters

x

Contents

11.5

12

Microfabricated Conjugated Polymer Actuators for Microvalves, Cell Biology, and Microrobotics Elisabeth Smela 12.1 12.2 12.3 12.4

12.5

12.6

12.7 12.8 12.9

13

14

11.4.3 Conjugated Polymer Based Steerable Catheters Discussion and Conclusion References

Introduction Actuator Background Microfabrication Single Hinge Bilayer Devices: Flaps and Lids 12.4.1 Bilayer Actuators 12.4.2 Drug Delivery 12.4.3 Cell Manipulation 12.4.4 Cell-Based Sensors Multi-Bilayer Devices: Positioning Tools 12.5.1 Microtools 12.5.2 Microrobot Swelling Film Devices: Valves 12.6.1 Out-of-Plane Actuation Strain 12.6.2 Microvalve Lifetime Integrated Systems Conclusions References

237 246 246

249 249 250 251 253 254 254 255 256 257 257 257 258 259 259 260 260 261 261

Actuated Pins for Braille Displays Geoffrey M. Spinks and Gordon G. Wallace

265

13.1 13.2 13.3 13.4 13.5 13.6

265 266 268 271 274 275 276 276

Introduction Requirements for the Electronic Braille Screen Mechanical Analysis of Actuators Operating Against Springs Polypyrrole Actuators for Electronic Braille Pins Other Polymer Actuation Systems for Electronic Braille Pins Summary Acknowledgements References

Nanostructured Conducting Polymer Biomaterials and Their Applications in Controlled Drug Delivery Mohammad Reza Abidian and David C. Martin 14.1 14.2

Introduction Nanostructured Conducting Polymers 14.2.1 Fabrication 14.2.2 Biomedical Application

279 279 280 280 282

Contents

14.3

14.4

15

Conducting Polymer Nanotubes for Controlled Drug Delivery 14.3.1 Electrospinning 14.3.2 Electrospinning of Dexamethasone-Loaded Template PLGA Nanofibers 14.3.3 Electrochemical Deposition of PEDOT Nanotubes 14.3.4 Controlled Drug Delivery from PEDOT Nanotubes Conclusions Acknowledgements References

285 286 287 288 289 293 293 293

Integrated Oral Drug Delivery System with Valve Based on Polypyrrole Thorsten Go¨ttsche and Stefan Haeberle

301

15.1 15.2 15.3

301 303 305 305 306 307 307 308 310 310 311 314 315 316

15.4

15.5

15.6

Introduction System Concept Osmotic Pressure Pump 15.3.1 Valve Closed 15.3.2 Valve Open Polypyrrole in Actuator Applications 15.4.1 Why PPy in the IntelliDrug System 15.4.2 Actuation of PPy Valve Concepts Evaluated in the Course of the IntelliDrug Project 15.5.1 Wafer-Level Fabricated Membrane Valve 15.5.2 Micro-Assembled Membrane Valve Total Assembly and Clinical Testing of the IntelliDrug System Acknowledgement References

SECTION IV 16

xi

PIEZOELECTRIC AND ELECTROSTRICTIVE POLYMERS 317

Piezoelectric and Electrostrictive Polymer Actuators: Fundamentals Zhimin Li and Zhongyang Cheng

319

16.1 16.2

319 320 320 321 323 324 325 326 326 328 328 330

16.3

16.4

Introduction Fundamentals of Electromechanical Materials 16.2.1 Piezoelectric Effect 16.2.2 Electrostrictive Effect 16.2.3 Other Effects Material Properties Related to Electromechanical Applications 16.3.1 Electromechanical Coupling Factor (k) 16.3.2 Elastic Response 16.3.3 Frequency and Temperature Responses Typical Electromechanical Polymers and Their Properties 16.4.1 Piezoelectric Polymers 16.4.2 Electrostrictive Polymers

xii

Contents

16.5

17

Miniature High Frequency Focused Ultrasonic Transducers for Minimally Invasive Imaging Procedures Aaron Fleischman, Sushma Srivanas, Chaitanya Chandrana and Shuvo Roy 17.1 17.2

17.3 17.4 17.5 17.6 17.7

18

Introduction Coronary Imaging Needs 17.2.1 Vulnerable Plaques 17.2.2 Stent Thrombosis High Resolution Ultrasonic Transducers 17.3.1 Polymer Transducers Fabrication Techniques Testing Methods Results Conclusion References

Catheters for Thrombosis Sample Exfoliation in Blood Vessels Using Piezoelectric Polymer Fibers Yoshiro Tajitsu 18.1 18.2 18.3 18.4 18.5

18.6

19

16.4.3 Maxwell Stress Effect Based Polymers 16.4.4 Practical Considerations Conclusions References

Introduction Piezoelectricity of Polymer Film and Fiber Simple Measurement Method for the Bending Motion of Piezoelectric Polymer Fiber Piezoelectric Motion of Poly-L-Lactic Acid (PLLA) Fiber Elementary Demonstration of Prototype System for Catheters Using Piezoelectric Polymer Fiber 18.5.1 Preliminary Demonstration 18.5.2 More Realistic Model for Application of Piezoelectric Polymer Fiber to Catheter Summary References

332 332 332 332

335

335 337 337 339 340 341 342 345 346 351 351

357 357 358 361 362 363 364 364 367 367

Piezoelectric Poly(Vinylidene) Fluoride (PVDF) in Biomedical Ultrasound Exposimetry Gerald R. Harris

369

19.1 19.2 19.3 19.4

369 370 371 372

Introduction Needle Hydrophone Design Spot Poled Membrane Hydrophone Design Application to Diagnostic Ultrasound

Contents

19.5 19.6

Application to Therapeutic Ultrasound Conclusion References

SECTION V 20

21

22

DIELECTRIC ELASTOMERS

xiii

374 377 378 385

Dielectric Elastomer Actuators: Fundamentals Roy Kornbluh, Richard Heydt and Ron Pelrine

387

20.1 20.2 20.3 20.4 20.5

387 388 389 391 392 393

Introduction Basic Principle of Operation Dielectric Elastomer Materials Transducer Designs and Configurations Operational Considerations References

Biomedical Applications of Dielectric Elastomer Actuators John S. Bashkin, Roy Kornbluh, Harsha Prahlad and Annjoe Wong-Foy

395

21.1 21.2 21.3 21.4 21.5 21.6 21.7 21.8

395 396 400 403 405 406 408 409 410

Introduction UMA Based Actuators and Their Application to Pumps Mechanical Stimulation Using Thickness-Mode Actuation Implantable Artificial Diaphragm Muscle Implantable Artificial Facial Muscles Limb Prosthetics and Orthotics Mechanical Actuation for ‘Active’ Cell Culture Assays Conclusions References

MRI Compatible Device for Robotic Assisted Interventions to Prostate Cancer Jean-Se´bastien Plante, Lauren Devita, Kenjiro Tadakuma and Steven Dubowsky 22.1 22.2

22.3

22.4

Introduction Prostate Cancer Therapy 22.2.1 Prostate Cancer Detection 22.2.2 Prostate Cancer Treatment 22.2.3 Needle Placement in MRI Systems Elastically Averaged Parallel Manipulator Using Dielectric Elastomer Actuators 22.3.1 Design Requirements 22.3.2 Manipulator Concept 22.3.3 Manipulator Analytical Model Results 22.4.1 Analytical Results 22.4.2 Experimental Results

411

411 413 413 414 415 415 415 417 418 420 421 423

xiv

Contents

22.5

23

A Braille Display System for the Visually Disabled Using a Polymer Based Soft Actuator Hyouk Ryeol Choi, Ig Mo Koo, Kwangmok Jung, Se-gon Roh, Ja Choon Koo, Jae-do Nam and Young Kwan Lee 23.1 23.2 23.3 23.4

23.5

23.6

24

Conclusions Acknowledgements References

Introduction Fundamentals of Actuation Principle Design of Tactile Display Device Braille Display System 23.4.1 Fabrication 23.4.2 System Outline 23.4.3 Experiments Advanced Applications 23.5.1 Wearable Tactile Display System 23.5.2 Virtual Reality Tactile Display Conclusions References

424 424 424

427

427 428 430 431 431 432 434 437 437 440 441 441

Dynamic Splint-Like Hand Orthosis for Finger Rehabilitation Federico Carpi, Andrea Mannini and Danilo De Rossi

443

24.1 24.2 24.3 24.4

443 444 445

Introduction Passive Dynamic Hand Splints: State of the Art Active Dynamic Hand Splints: State of the Art Proposed Concept: Dynamic Splint Equipped with Dielectric Elastomer Actuators 24.5 Splint Mechanics 24.6 Dimensioning of the Actuators 24.7 Prototype Splint 24.8 Performance of the Prototype Splint 24.9 Future Developments 24.9.1 Magnetic Resonance Imaging-Compatible Hand Splint 24.9.2 Electromyography-Controlled Hand Splint 24.10 Conclusions References Index

446 449 449 450 451 454 454 457 460 460 463

Preface

The great majority of traditional actuation technologies are based on thermochemical (combustion) motors, electromagnetic drives and hydraulic/pneumatic machines. However, these are inadequate to satisfy the diversity of new challenges presented in fields such as mechatronics, robotics and biomedical engineering. The biomedical field is particularly sensitive to the need for new types of actuators, since it includes applications from the nanoscale through to the macroscale, with requirements that differ enormously in terms of both structure and function. As an example, actuation devices span those designed to interact with single cells, and potentially organelles and even molecular structures, up to those used to replace limbs or to perform tele-operated surgery. Clearly, these different areas of application require different forces, displacements and speeds, as well as different durability, robustness and types of biocompatibility. Devices for cell manipulation might be single-use and disposable, whereas artificial hearts must sustain billions of cycles. New actuation materials and technologies should ideally have high work output, actuation strain, mechanical compliance, damage tolerance and efficiency. Depending on the application, they may also be required to be lightweight, have compact and simple structures that can readily be fabricated, and be reasonably low cost. In most scenarios, these new technologies will serve a complementary role working together with conventional actuators. In the last few decades this need for new actuators has drawn considerable effort towards the development of materials that can directly transduce an input energy into mechanical work. Much of this attention has been focused, and is increasingly being focused today, on electroactive polymers (EAPs). This is a large family of materials that includes many different chemical structures, actuation mechanisms and electromechanical performances. Although most EAP materials have been known for decades, before now they found limited translation from proof of concept demonstrations in the laboratory to actual use, despite their potential. This has changed with recent developments in materials science, processing, configuration design and driving strategies which are permitting serious efforts towards concrete exploitation, as this book describes. In fact, EAPs are opening the way to numerous new applications precluded by conventional actuation technologies. This book intends to provide a comprehensive and updated insight into both the fundamentals of each class of EAP, and examples of the most significant applications of EAP actuators in the biomedical field, either already demonstrated or currently under development. For this purpose, the book comprises five sections devoted to the most technologically mature EAPs, namely polymer gels, ionic polymer–metal composites, conjugated polymers, piezoelectric/electrostrictive polymers and dielectric elastomers. Each section is

xvi

Preface

introduced by a chapter that is focused on the fundamentals and which aims to provide a description of the main features of the technology and the current state of the art. These introductory chapters are followed by chapters describing specific applications. The contributors to this book are inventors and international leaders in the field. The broad and far-reaching range of applications covered by this book is intended not only to make it the first text on biomedical uses of the emerging EAP based actuation technologies, but also to serve as a source of inspiration for possible new applications aimed at improving health and well-being. Federico Carpi, University of Pisa Elisabeth Smela, University of Maryland October 2008

List of Contributors

Cameron Alexander, School of Pharmacy, University of Nottingham, UK Gursel Alici, ARC Centre of Excellence for Electromaterials Science and Intelligent Polymer Research Institute, University of Wollongong, Australia Kinji Asaka, National Institute of Advanced Industrial Science and Technology (AIST), Japan Mark Banister, Medipacs LLC, Tucson, USA John Bashkin, Fremont, CA, USA Paul Calvert, University of Massachusetts, Dartmouth, USA Federico Carpi, Interdepartmental Research Centre ‘‘E. Piaggio’’, School of Engineering, University of Pisa, Italy Chaitanya Chandrana, Cleveland Clinic, Lerner Research Institute, Department of Biomedical Engineering, Cleveland, USA Zhongyang Cheng, Materials Research and Education Center, Alkermes Inc., Auburn, Alabama, USA Piero Chiarelli, Institute of Clinical Physiology, CNR, Italy Hyouk Ryeol Choi, School of Mechanical Engineering, Chemical Engineering, Polymer System Engineering, Sungkyunkwan University, Korea Brian G. Cousins, School of Materials, Materials Science Centre and Manchester Interdisciplinary Biocentre (MIB), University of Manchester, United Kingdom Danilo De Rossi, Interdepartmental Research Centre ‘‘E. Piaggio’’, School of Engineering, University of Pisa, Italy Lauren Devita, Massachusetts Institute of Technology, USA

xviii

List of Contributors

Steven Dubowsky, Massachusetts Institute of Technology, USA Aaron Fleischman, Cleveland Clinic, Lerner Research Institute, Department of Biomedical Engineering, Cleveland, USA Thorsten Go¨ttsche, Institut fu¨r Mikro- und Informationstechnik of the Hahn-SchickardGesellschaft (HSG-IMIT), Germany Stefan Haeberle, Institut fu¨r Mikro- und Informationstechnik of the Hahn-SchickardGesellschaft (HSG-IMIT), Germany Gerald R. Harris, Food and Drug Administration, Center for Devices and Radiological Health, USA Richard Heydt, SRI International, USA Kwangmok Jung, Pohang Institute of Intelligent Robotics, Korea Kwang J. Kim, University of Nevada, USA Sang-Mun Kim, University of Nevada, USA Masashi Konyo, Graduate School of Information Sciences, Tohoku University, Japan Ig Mo Koo, School of Mechanical Engineering, Chemical Engineering, Polymer System Engineering, Sungkyunkwan University, Korea Ja Choon Koo, School of Mechanical Engineering, Chemical Engineering, Polymer System Engineering, Sungkyunkwan University, Korea Roy Kornbluh, SRI International, USA Sangki Lee, University of Nevada, USA and Volvo Korea, South Korea Young Kwan Lee, School of Mechanical Engineering, Chemical Engineering, Polymer System Engineering, Sungkyunkwan University, Korea Zhimin Li, Pharmaceutical Chemistry, Auburn University, Cambridge, Massachusetts, USA John D. Madden, Advanced Materials and Process Engineering Laboratory and Department of Electrical & Computer Engineering, University of British Columbia, Vancouver, Canada Jo´hannes Pa´ll Magnu´sson, School of Pharmacy, University of Nottingham, UK Andrea Mannini, Interdepartmental Research Centre ‘‘E. Piaggio’’, School of Engineering, University of Pisa, Italy

List of Contributors

xix

David C. Martin, Biomedical Engineering, Materials Science and Engineering and Macromolecular Science and Engineering, The University of Michigan, Ann Arbor, MI, USA Tom McDonald, School of Materials, Materials Science Centre, University of Manchester and Manchester Interdisciplinary Biocentre (MIB), University of Manchester, United Kingdom Scott McGovern, ARC Centre of Excellence for Electromaterials Science and Intelligent Polymer Research Institute, University of Wollongong, Australia Nigel R. Munce, Imaging Research, Sunnybrook Health Science Centre, University of Toronto, Canada Jae-do Nam, School of Mechanical Engineering, Chemical Engineering, Polymer System Engineering, Sungkyunkwan University, Korea Keisuke Oguro, National Institute of Advanced Industrial Science and Technology (AIST), Japan Il-Seok Park, University of Nevada, USA Alison Patrick, School of Materials, Materials Science Centre and Manchester Interdisciplinary Biocentre (MIB), University of Manchester, United Kingdom Ron Pelrine, SRI International, USA Jean-Se´bastien Plante, Universite´ de Sherbrooke, Canada Harsha Prahlad, SRI International, USA Pietro Ragni, Institute of Nuclear Chemistry, CNR, Italy Mohammad Reza Abidian, Biomedical Engineering, The University of Michigan, Ann Arbor, MI, USA Se-gon Roh, School of Mechanical Engineering, Chemical Engineering, Polymer System Engineering, Sungkyunkwan University, Korea Shuvo Roy, Cleveland Clinic, Lerner Research Institute, Department of Biomedical Engineering, Cleveland, USA Aram Omer Saeed, School of Pharmacy, University of Nottingham, UK Mohsen Shahinpoor, Biomedical Engineering Laboratories, Department of Mechanical Engineering, University of Maine, Orono, USA

xx

List of Contributors

Tina Shoa, Advanced Materials and Process Engineering Laboratory and Department of Electrical & Computer Engineering, University of British Columbia, Vancouver, Canada Elisabeth Smela, Department of Mechanical Engineering, University of Maryland, USA Geoffrey M. Spinks, ARC Centre of Excellence for Electromaterials Science and Intelligent Polymer Research Institute, University of Wollongong, Australia Sushma Srivanas, Cleveland Clinic, Lerner Research Institute, Department of Biomedical Engineering, Cleveland, USA Kenjiro Tadakuma, Massachusetts Institute of Technology, USA Satoshi Tadokoro, Graduate School of Information Sciences, Tohoku University, Japan Yoshiro Tajitsu, Smart Structures and Materials Laboratory, Department of Electrical Engineering, Graduate School of Engineering, Kansai University, Japan Beverley Twaites, School of Pharmacy and Biomedical Sciences, University of Portsmouth, UK Rein V. Ulijn, University of Strathclyde, United Kingdom Sonia Vohnout, Medipacs LLC, Tucson, USA Gordon G. Wallace, ARC Centre of Excellence for Electromaterials Science and Intelligent Polymer Research Institute, University of Wollongong, Australia Richard Williams, School of Materials, Materials Science Centre and Manchester Interdisciplinary Biocentre (MIB), University of Manchester, United Kingdom Annjoe Wong-Foy, SRI International, USA Binbin Xi, ARC Centre of Excellence for Electromaterials Science and Intelligent Polymer Research Institute, University of Wollongong, Australia Victor X.D. Yang, Imaging Research, Sunnybrook Health Science Centre and Department of Electrical and Computer Engineering, Ryerson University, Toronto, Canada

Plate 1

A heart with an IPMNC compression band. (See Figure 7.17)

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

2.58 10 5

25

10 63 Laser beam line width: 0.2 (a)

(mm)

(b)

Plate 2 (a) Laser cutting machine and (b) a CAD design for interdigitated IPMC (Reproduced with permission from Vohnout, S., Kim, S.-M., Park, I.-S. and Banister, M., IPMC-assisted miniature disposable infusion pumps with embedded computer control, Proceedings of the SPIE conference 2007. Copyright (2007) SPIE). (See Figure 9.5)

MSC.Patran 2001 r2a 16-Aug04 10:12:28 Fringe: SC1:DIAPHRAGM, A14:Static Subcase: Displacements, Translational-(NON-LAYERED) (ZZ) Deform: SC1:DIAPHRAGM, A14:Static Subcase: Displacements, Translational

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9.66–004

0.

4.51–004 3.87–004

+

3.22–004 2.58–004 1.93–004 1.29–004 6.44–005

Z

–1.16–010 Y X

default_Fringe: Max 9.66–004 @Nd1 Min 0. @Nd 316 default_Deformation : Max 9.66–004 @Nd1

(a)

MSC.Patran 2001 r2a 16-Aug04 10:30:37 Fringe: SC1:DIAPHRAGM, A11:Static Subcase: Displacements, Translational-(NON-LAYERED) (ZZ) Deform: SC1:DIAPHRAGM, A11:Static Subcase: Displacements, Translational

0 –4.57–005 –9.14–005 –1.37–004 –1.83–004 –2.29–004 –2.74–004 –3.20–004

0. +

–3.66–004 –4.11–004 –4.57–004

–0

–5.03–004 –5.48–004 –5.94–004 –6.40–004 Z

–6.86–004 Y X

(b)

default_Fringe: Max 0. @Nd316 Min –6.86–004@Nd1 default_Deformation : Max 6.86–004 @Nd1

Plate 3 Deformed shapes of IPMC diaphragms: (a) circle-shaped electrode (radius of electrode ¼ 8.5 mm); (b) ring-shaped electrode (radial length of electrode ¼ 5.5 mm) (Reproduced with permission from Lee, S., Kim, K.J. and Park, H.C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications). (See Figure 9.12)

MSC.Patran 2001 r2a 20-Aug04 14:27:41 Fringe: SC1:DIAPHRAGM, A2:Mode 1: Freq. = 429.69: Eigenvectors, Translational-(NON-LAYERED) (MAG) Deform: SC1:DIAPHRAGM, A2:Mode 1: Freq. = 429.69: Eigenvectors, Translational

3.15+002 2.94+002 2.73+002 2.52+002 2.31+002 2.10+002 1.89+002 1.68+002

3.15+002

1.47+002

0.

1.26+002

+

1.05+002 8.41+001 6.31+001 4.21+001 2.10+001

Z

4.20–005 Y X

default_Fringe: Max 3.15+002 @Nd1 Min 0. @Nd 316 default_Deformation : Max 3.15+002 @Nd1

(a)

MSC.Patran 2001 r2a 20-Aug-04 14:28:51 Fringe: SC1:DIAPHRAGM, A2:Mode 2: Freq. = 1659.1: Eigenvectors, Translational-(NON-LAYERED) (MAG) Deform: SC1:DIAPHRAGM, A2:Mode 2: Freq. = 1659.1: Eigenvectors, Translational

3.02+002 2.81+002 2.61+002 2.41+002 2.21+002 2.01+002 1.81+002 1.61+002

0 3.02+0

2

0.

1.41+002 1.21+002

+

1.01+002 8.04+001 6.03+001 4.02+001 2.01+001

Z

3.81–006 Y X

(b)

default_Fringe: Max 3.02+002 @Nd261 Min 0. @Nd316 default_Deformation : Max 3.02+002 @Nd261

Plate 4 Normal mode analysis results for an IPMC diaphragm (radius of electrode ¼ 8.5 mm): (a) first mode; (b) second mode (Reproduced with permission from Lee, S., Kim, K.J. and Park, H.C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications). (See Figure 9.14)

Plate 5 Schematic diagrams illustrating the surface modification of neural microelectrodes to create nanotubular PEDOT: (A) electrospinning of biodegradable polymer (PLGA) fibers with well-defined surface texture (1) on the probe tip; (B) electrochemical polymerization of conducting polymers (PEDOT) (2) around the electrospun fibers; and (C) dissolving the electrospun core fibers to create nanotubular conducting polymers (3) [7] (Reprinted wih permission from Advanced Materials, Conducting polymer nanotubes for controlled drug release by Abidian, M. R., et al., 18, 4, 405–9. Copyright (2006) Wiley-VCH Verlag GmbH Co. KGaA). (See Figure 14.5)

Plate 6 Optical micrographs of: (E) the gold electrode site; (F) the electrode site after electrospinning showing the coverage of the PLGA electrospun nanoscale fibers; (G) the electrode after electrochemical deposition of PEDOT on the gold site and around the electrospun fibers; and (H) the electrode after removal of the core nanoscale fiber templates (Reprinted with permission from Advanced Materials, Conducting polymer nanotubes for controlled drug release by Abidian, M. R., et al., 18, 4, 4059. Copyright (2006) Wiley-VCH Verlag GmbH Co. KGaA). (See Figure 14.7)

A

C

B

D

E

F CE

CE

v

v

WE

WE

Electrolyte

Dexamethasone Anion Cation

CE Counter Electrode WE Working Electrode

Plate 7 Schematic illustration of the controlled release of dexamethasone: (A) dexamethasoneloaded electrospun PLGA; (B) hydrolytic degradation of PLGA fibers leading to release of the drug; and (C) electrochemical deposition of PEDOT around the dexamethasone-loaded electrospun PLGA fiber slows down the release of dexamethasone (D); (E) PEDOT nanotubes in a neutral electrical condition; (F) external electrical stimulation controls the release of dexamethasone from the PEDOT nanotubes due to contraction or expansion of the PEDOT (Reprinted with permission from Advanced Materials, Conducting polymer nanotubes for controlled drug release by Abidian, M. R., et al., 18, 4, 405–9. Copyright (2006) Wiley-VCH Verlag GmbH Co. KGaA). (See Figure 14.9)

1.0 Absorbance (AU)

Cumulative Mass Released (mg)

2.0 1.5 1.0 0.5 0.0

0

200 400 600 800 1000 1200 1400 Time (h) (a)

0.8 0.6 0.4 0.2 0.0

220 240 260 280 300 Wavelength (nm)

320

(b)

Plate 8 (a) Cumulative mass release of dexamethasone from: PLGA nanoscale fibers (black squares), PEDOT-coated PLGA nanoscale fibers (red circles) without electrical stimulation, and PEDOT-coated PLGA nanoscale fibers with electrical stimulation of 1 V applied at the five specific times indicated by the circled data points (blue triangles). (b) UV absorption of dexamethasone- loaded PEDOT nanotubes after 16 h (black), 87 h (red), 160 h (blue) and 730 h (green). The UV spectra of dexamethasone have peaks at a wavelength of 237 nm. Data are shown with a  standard deviation (n ¼ 15 for each case) (Reprinted with permission from Advanced Materials, Conducting polymer nanotubes for controlled drug release by Abidian M.R., et al., 18, 4, 405–9. Copyright (2006) Wiley-VCH Verlag GmbH Co. KGaA). (See Figure 14.10(a) and 14.10(b))

Introduction Electroactive Polymers as Smart Materials for Actuation Federico Carpi 1 and Elisabeth Smela 2 1

Interdepartmental Research Centre ‘E. Piaggio’, University of Pisa, Italy 2 Department of Mechanical Engineering, University of Maryland, USA

I.1 Actuation: the Need for New Materials and Technologies Actuators are materials, devices or systems that are able to act upon their external environment by transducing input energy into external mechanical work. Biological muscles have long drawn interest as natural actuation systems, and many regard their functional properties as ideal models for artificial actuators. The conventional actuators used today, which mainly comprise thermochemical motors (combustion engines), electromagnetic drives and hydraulic/pneumatic machines, differ considerably from their natural counterparts. For example, their underlying physical principles of actuation share little with biological muscles. In fact, the latter are electro-chemo-mechanical actuators based on contractile proteins that use the body’s chemical energy (adenosine triphosphate, ATP) to generate motion, triggered by neuro-electric commands. Despite the advanced performance of conventional motor technologies, which in some respects outreaches that of natural muscles, they cannot satisfy the ever-increasing demand for new actuators in a large number of quite different areas of technology, spanning the range from mechatronics to biomedical engineering. In particular, new actuation technologies are needed that are easily scalable, structurally simple and mechanically compliant, while also having high power-to-weight and power-to-volume ratios, and fine control capability. To illustrate the need for new approaches to actuation, consider that despite years of effort to develop prostheses (artificial hands and arms) driven by electric motors, currently available systems are still stiff, heavy and noisy. The effort to develop ‘artificial muscles’

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Biomedical Applications of Electroactive Polymer Actuators

would benefit from new actuators with intrinsic properties similar to those of biological muscular tissue, in terms of both passive characteristics (namely high mechanical compliance and low mass density) and active behaviour (intrinsic actuation capabilities in response to electrical stimuli). Another consideration is the small size of the objects that it is wished to move or manipulate, for which miniaturization of conventional actuators is not going to be possible. For instance, insect-size robots that can fly, hop and crawl are one example of systems that cannot be built with conventional actuators. Fabrication paradigms and driving principles prevent conventional drives from being scaled down further. The inability of traditional actuation technologies to cover the entire spectrum of requirements for new, and vastly different, application areas has given rise to new materials, new drive principles and new devices. These are intended to play at least a complementary role, so as to compensate the deficits of the more consolidated technologies in specific domains. To satisfy the needs described above, new research avenues focused on electroactive polymers have been opened in recent years.

I.2 Electroactive Polymers: Classification Polymers are promising candidates for new actuation technologies. Several classes of polymeric compounds can convert electrical energy (or other sources of energy, such as heat, light and chemical gradients) into mechanical work, so as to effect the movement of loads. Piezoelectric and shape memory polymers are among the most well known representatives of polymer actuators. Electroactive polymers (EAPs) [1, 2] comprise a broad family not only of active materials like piezoelectrics that intrinsically change volume, but also of polymercontaining devices in which the polymer is passive. They share the capability of changing dimensions and/or shape in response to a stimulus, which is most preferably electrical. They can have sizable actuation strains compared with inorganic materials (although actuation stresses are lower), high compliance, low mass density and scalability. They also promise ease of processing and low cost. As a result, EAPs are one of the most promising classes of materials for ‘muscle-like’ actuators [1, 2]. EAPs can be divided into two main classes: ionic, which are activated by electricallyinduced transport of ions and/or solvent, and electronic, which are activated by electric fields [1, 2]. The former generally require low voltage but high currents, while the latter need high voltage but low currents. They thus have different application areas, as described in detail in the later chapters of this book. Another difference is that the ionic EAPs operate in a liquid electrolyte medium, while the electronic EAPs are used in air. The ionic EAPs are thus particularly attractive for use within biological environments. The electronic EAPs can operate at higher frequencies and with higher efficiencies in converting electrical energy into mechanical work, making them of more interest in applications such as robotics. No EAP technology can today play a general-purpose role, and the selection of the most suitable actuator should be carefully evaluated according to the specific requirements of the application. Within each group there are a wide variety of specific actuation mechanisms and related types of materials. In particular, ionic EAPs include polymer gels [3], ionic polymer–metal

Electroactive Polymers as Smart Materials for Actuation

3

composites (IPMC) [4], conjugated (conducting) polymers [5] and carbon nanotubes [6]. Electronic EAPs include piezoelectric polymers [7], electrostrictive polymers [8], flexoelectric polymers [9] and dielectric elastomers [10]. The different types of EAPs and the most significant materials used in each type are summarized in Table I.1. Table I.1

Electroactive polymer actuators: classification and representative materials

EAP class

EAP sub-categories

Ionic EAP

Polymer gels

Electronic EAP

Examples of materials

Poly(acrylic acid) (PAAc) Poly(vinyl alcohol) (PVA) Modified poly(acrylonitrile) (PAN) Ionic polymer–metal Metalized ion exchange membranes e.g. composites (IPMC) Nafion/Pt Conjugated polymers Polypyrrole (PPy) Polyaniline (PANi) Carbon nanotubes Single-walled nanotubes (SWCNT) Multi-walled nanotubes (MWCNT) Piezoelectric polymers Poly(vinylidene fluoride) (PVDF) Electrostrictive polymers PVDF based copolymers e.g.: Poly(vinylidene fluoride–trifluoroethylene) (PVDF–TrFE) Poly(vinylidene fluoride– hexafluoropropylene) (PVDF–HFP) Flexoelectric polymers Liquid crystal elastomers Dielectric elastomers Silicone elastomers Acrylic elastomers Polyurethane elastomers

Ref. [3] [4] [5] [6] [7] [8]

[9] [10]

The remainder of this book is organized into sections covering each of the main classes of EAPs. The first chapter of each section focuses on the fundamentals of each technology, describing the actuation mechanism and performance. Two of the classes listed in Table I.1, carbon nanotubes and flexoelectric polymer actuators, have not been included in this book because of their lower technological maturity.

References 1. Bar-Cohen, Y. (2004) (ed.), Electroactive Polymer (EAP) Actuators as Artificial Muscles: Reality, Potential, and Challenges, 2nd edn, SPIE Press Monograph, Vol. PM136. 2. Madden, J. D. W., Vandesteeg, N. A., Anquetil, P. A., et al. (2004) Artificial muscle technology: physical principles and naval prospects, IEEE J. Oceanic Eng., 29, 3, 706–28. 3. Tanaka, T., Nishio, I., Sun, S-T. and Ueno-Nishio, S. (1982) Collapse of gels in an electric field, Sci., 218, 467–9. 4. Asaka, K., Oguro, K., Nishimura, Y., et al. (1995) Bending of Polyelectrolyte MembranePlatinum Composites by Electric Stimuli, I. Response Characteristics to Various Waveforms, Polym. J. 27, 4, 436–40. 5. Baughman, R. H. (1996) Conducting polymer artificial muscles, Synth. Met., 78, 339–53. 6. Baughman, R. H., Changxing, C., Zakhidov, A. A., et al. (1999) Carbon nanotube actuators, Sci., 284, 1340.

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Biomedical Applications of Electroactive Polymer Actuators

7. Nalwa, H. S. (1995) Ferroelectric Polymers, Marcel Dekker, New York. 8. Zhang, Q. M., Bharti, V., Zhao, X. (1998) Giant electrostriction and relaxor ferroelectric behaviour in electron-irradiated poly(vinylidene fluoride–trifluoroethylene) copolymer, Sci., 280, 2101–3. 9. Lehmann, W., Skupin, H., Tolksdorf, C., et al. (2001) Giant lateral electrostriction in ferroelectric liquid-crystalline elastomers, Nat., 410, 447–50. 10. Pelrine, R., Kornbluh, R., Pei, Q. and Joseph, J. (2000) High-speed electrically actuated elastomers with strain greater than 100%, Sci., 287, 836–839.

Section I Polymer Gels

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

1 Polymer Gel Actuators: Fundamentals Paul Calvert University of Massachusetts, Dartmouth, USA

1.1 Introduction and Historical Overview One view of materials development is to search for what materials correspond to empty spaces on a hypothetical multi-dimensional map of the properties of available materials. Following a biomimetic philosophy, for instance, it can be seen that tough ceramics and moldable short fiber composites with high moduli are possible but absent from the list of available synthetic materials. Likewise artificial muscle is missing, where the properties are defined as a developed stress of over 300 kPa, a linear contraction of 25 % and a response time of below one second. Currently dielectric elastomers come closest but have disadvantages [1, 2]. The actin–myosin muscle system provides the performance target for electroactive polymer actuators [3, 4]. The process is driven chemically by the energy change from hydrolysis of the polyphosphate bond as ATP (adenosine triphosphate) binds to myosin and is converted to ADP (adenosine diphosphate) and phosphate. A simple chemical analogy suggests that muscle-like gels should be feasible but it is now clear that the task is much harder than it seems. Early work on gel actuation by Katchalsky–Katzir demonstrated that engines could be built using the chemical energy in diluting a lithium bromide solution to drive contraction and expansion of a collagen belt [5]. In essence, the collagen expands to take up the salt solution or contract to exclude the water. This change is both a molecular conformation change and a volume change. Other chemically driven gels, for instance polyacrylic acid fibers [6] which respond to a pH change, also rely on a solubility change giving rise to a volume change.

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Biomedical Applications of Electroactive Polymer Actuators

Work by Osada and others explored electrically-driven actuators where pH changes at each electrode drive local volume changes [7]. While electrical power should be a more convenient way of driving artificial muscles, it has proved to be difficult to couple the electrical energy input to provide a mechanical output. As a result the achievable efficiencies for energy conversion are very low. These systems driven by volume changes are rate-limited by the need for water or solvent to diffuse in and out of the gel. The force developed is also limited by the free energy associated with the change in solvent content. In this regard, the apparent analogy between muscle action and gel contraction breaks down. Muscle contraction is essentially a shape change driven by a conformation change. As such there seems to be no significant volume change and so no diffusion limitation. The energy source is dephosphorylation of ATP, which is also large compared to many solvation interactions. The fact that early work of Katchalsky used collagen, which so resembles muscle, may have been misleading. Subsequent work that has focused on similar artificial muscles has developed systems that are intrinsically too slow and too feeble to provide the force and speed associated with natural muscle. In this chapter the properties and applications of current gel actuators are summarized and the developments needed to produce a true artificial muscle are discussed.

1.2 Properties of Gels To be useful as actuators, gels must be stiff enough to exert the desired force and must be strong enough to carry the desired load. If the change in shape is caused by swelling, the fluid must flow into the gel in a short time. The transport of fluid will also be important to whether the actuator can be used without drying out. In addition, because gels cannot be readily machined or molded, it must be possible to form the gel into a suitable shape for an actuator by polymerization in situ. If gels are to be used as actuators, they must be mechanically robust. Most current applications of gels are in food, pharmaceuticals and cosmetics where mechanical properties are not a major concern. Muscle as an actuator develops a force of about 300 kPa and a strain of 25 %. If these are regarded as target numbers, a gel actuator should be strong enough to easily withstand the force developed, and so the fracture strength should be at least 1 MPa. In addition, it should be stiff enough that, in lifting a weight for instance, an active contraction is not simply cancelled by passive extension under the load. If it is said that the passive extension should be less than one-tenth of the contraction, the elastic modulus would be greater than 12 MPa. While a low force gel actuator may be useful for some applications, such as a microfluidics valve, the mechanical requirements on a muscle-like actuator are quite demanding. 1.2.1

Biological Gels

In biology there are many examples of structural gels in the marine environment, including seaweeds and the bodies of many invertebrates, such as sea anemones. In the human body, cartilage, cornea, the dermis and arterial wall are all fiber-reinforced gels. Although soft and not very strong, these materials are very tough and so survive the impacts of life in motion better than many hard materials in machines.

Polymer Gel Actuators: Fundamentals

9

Many biological tissues show a ‘J-shaped’ tensile stress–strain curve [8] that combines a low initial modulus with high strength. The reinforcing fibers initially rotate as the soft gel is stretched and then the fibers take up the load as they become parallel to the stress axis. Other variations occur depending on whether the fibers are bonded into a network. Unbonded fibers can flow with the gel under slow loading but give rigidity under fast loading [9]. Articular cartilage is a proteoglycan gel reinforced with about 20 % of collagen fibers. It has a strength of about 1.5 MPa and an extension to break of about 100 %. The structure is layered and the properties vary greatly with depth below the surface, with strength up to up to 30 MPa in layers with higher fiber contents [8, 10–13]. Costal (rib) cartilage has a strength of 5–7 MPa [14]. The large extension to break and large work of fracture (1 kJ/m2) [15] allow cartilage to function effectively under impact even though the average strength is not high. An unfamiliar aspect of the mechanical properties of gels is that they will tend to lose water under compression and take up water under tension. As a result, the mechanical properties will be different depending on the test speed. Thus the fast modulus of cartilage is 2.5 MPa while the equilibrium modulus, measured as water is displaced from the structure, is about 0.7 MPa [8, 12]. The transition between these two values will depend on sample size, as the water has to flow out of the gel. Likewise, testing under water will result in properties that differ from those measured in air. Cornea is another tissue that is reinforced with collagen fibers. As with cartilage, there is an immediate need for a synthetic substitute to replace damaged corneal tissue. Cornea contains about 20 % of collagen fibers in a gel matrix. The fiber diameter is of the order of 20 nm, so that light is not scattered and the material is transparent [16]. The tensile strength is about 4 MPa and the elastic modulus is about 6 MPa at 20 % strain on the J-shaped stress–strain curve [17]. The properties of these soft tissues and marine gels suggests that a combination of a higher polymer content and fiber reinforcement should let us form materials which retain the responsive properties of gels whilst having sufficient mechanical strength to be used in engineering systems. As a target system, we could envision seeking a synthetic gel with a tensile strength of 5 MPa and an extension to break of 100 % but we do also need to develop a better understanding of the mechanical properties of gels. 1.2.2

Mechanical Properties of Simple, Single-Phase Gels

Because gels are weak and so do not currently have many synthetic applications, there is not a large coordinated literature on their mechanical properties. For many materials we can consider elastic modulus and tensile strength as sufficient to characterize the mechanical properties. The first of these reflects the rigidity or degree of bending under stress, the second the ability to withstand static stress without breaking or deforming irreversibly. For hard materials that have to withstand impacts, we are also concerned with the toughness, often measured as the energy absorbed in propagating a crack through the material. These concepts are substantially derived from the consideration of metals and ceramics where their stress–strain relationship is essentially linear up to about 1 % strain and then yield or fracture occurs. They serve well in most engineering situations where objects are designed to be rigid.

10

Biomedical Applications of Electroactive Polymer Actuators

Many tissues operate in a different regime, where there are large reversible strains and substantial impacts can be tolerated without damage. In this case the energy needed to produce damage may be more important than the strength. The stress–strain curve may be very nonlinear and the shape of the curve out to large strains becomes important. The same is true of rubbery materials, but these are also not very familiar in structural engineering. Thus, engineering with gels will put us into a regime that is unfamiliar to many mechanical engineers. 1.2.3

Elastic Moduli

One view of a gel is as a modified rubber. The properties of amorphous polymers change dramatically above the glass transition point where the chains become mobile. Since polymer chains are mobile in solution and we think of a simple gel as a cross-linked solution, we can regard a single phase gel as a dilute rubber. Dense polyacrylamide, for instance, is a glassy polymer. We do not want to compare the gel properties with this state but with the same polymer as a cross-linked rubber above its glass transition. Gels are soft materials, so we would expect elastic moduli to be below 10 MPa and we would expect the modulus to decrease as the volume fraction of solvent increases. As an example, a gelatin gel swollen to five times its dry weight, has a modulus of about 0.8 MPa and a fracture stress of about 70 kPa with an extension to break of 10 %. At a swelling of 40 times, the modulus is only 40 kPa and the strength 6 kPa with the extension to break 11 %. This soft, weak, brittle behavior is characteristic of most simple gels. The initial elastic modulus of a nonionic gel can be derived from rubber elasticity theory as: G¼A


2=3  RT v02 ðv2 Þ1=3 Mc

ð1:1Þ

where A is close to one, r is density, Mc is the average molecular weight between cross-links, v2 is the volume fraction of polymer in the swollen gel and v20 is the volume fraction in the gel as synthesized. Thus swelling after synthesis decreases the modulus relative to the Go, the modulus as synthesized [18, 19]:  1=3 G v2 ¼ ¼ Vr 1=3 ð1:2Þ G0 v02 Many natural gels are highly charged polyelectrolytes. It might be expected that there would be a strong difference in modulus between otherwise comparable ionic and nonionic gels. For weakly charged groups (acrylic acid), charge seems to have little effect on elastic properties [20]. Other studies show that charged groups increase the modulus and decrease the dependence of modulus on swelling [21]. 1.2.4

Strength

Synthetic gels, based on cross-linked soluble polymers, are mostly too weak to be used in any structural application. Many natural gel structures, such as are found in marine organisms, seem to be quite strong. As discussed next, possibly this difference arises from the microstructure of natural gels, which most synthetic gels lack.

Polymer Gel Actuators: Fundamentals

11

In determining the strength of gels, there are important factors which can often be ignored in dense materials. Gels often fracture at much higher strains than conventional engineering materials, properties can be very time dependent and liquid may be taken up or lost during the test. Likewise, the properties of immersed gel samples differ from samples tested in air, as water is normally taken up in tension and exuded in compression. The degree of confinement and timescale of testing is also going to be important for the same reasons. At high compressive strains sample geometry will also be crucial, since friction at the platens will result in shear stresses which effectively put the sample into hydrostatic compression where failure cannot occur. Thus properties at high strain in compression must not be regarded as directly comparable to a true strength. Most hard amorphous polymers under tension show brittle fracture. The strength, s, is determined under the Griffith equation (Equation (1.3)) by the fracture surface energy, g, which in turn mostly depends on the energy absorbed by the plastic deformation and void formation (crazing) that occurs immediately at the tip of the crack. E is the elastic modulus and c is the crack length:   2 E 1=2 ¼ ð1:3Þ pc Since elastomers are essentially liquid polymers, the elastic modulus is low and crazing is not believed to occur. Most of the fracture energy probably goes into pulling individual chains out across the fracture and so the energy increases with the chain length between cross-links [22]. Fracture of rubbers does not follow the Griffith theory because of the higher extensions at fracture, but the role of fracture energy in limiting crack extension still applies. Most unreinforced elastomers lack significant energy absorbing mechanisms and so readily tear at any cut or notch. Based on the comparison between gels and elastomers, we would expect the strength of unstructured gels to be lower than that of rubbers with a similar cross-link density by factors reflecting the dilution of the gel by water or solvent and reflecting the degree of preextension of the chain due to swelling by the solvent. We thus expect gels to be weak and to get weaker as they swell more. There are exceptions to the generally low strength of elastomers where some energyabsorbing deformation can occur. One example is natural rubber, where the crystallization occurs under tension, resulting in increased stiffness at high stress and a large energy to fracture as chains slip through the ordered crystals. Large fracture energies are also obtained when diene rubber chains slip over the surface of reinforcing carbon black particles or through the hard regions of two-phase polyurethane elastomers. It may be possible to build similar energy-absorbing mechanisms into gels. Theoretical discussion of gel fracture has focused on gelatin gels, which are important in food. Both fracture mechanics and fracture energy approaches have been considered but understanding is still imperfect [23–25]. Synthetic gels based on acrylates are unstructured and so would also be expected to be weak. Natural gels, such as agarose [26] and the calcium alginates, do seem to form ordered regions of double helix or multiple helices [27–29]. As a result, the mechanical properties are very dependent on the extent of structure developed during gelation [30]. The disruption of these structures during fracture could be expected to be a source of energy

12

Biomedical Applications of Electroactive Polymer Actuators

absorption and so increase strength and toughness. 2 % Agarose gels have a strength of about 0.14 MPa and a strain to failure of 40 %. In contrast, the strength of similar gelatin gels is about 1 kPa [31]. It is possible that similar ordered structures, to those in agarose, exist in gels of hydrogen-bonding polymers, such as hydroxyethylmethacrylate and vinylpyrrolidone, which do seem to be stronger than less polar synthetic gels. One area where there has been a vigorous search for improved mechanical properties is in gels for contact lenses but there is no clear picture of what determines strength [32, 33]. Contact lenses have water contents of 30–50 % and strengths of 2–4 MPa [34]. Tests on vinylpyrrolidone gels with low water contents (30–40 %) gave strengths up to 2 MPa, which is in the range of cartilage and so could be considered adequate for construction of equipment [35]. On the other hand, contact lens gels made from mixed acrylic and vinylpyrrolidone monomers with 40–70 % water content have strengths from 100–600 kPa [36]. There is no simple relationship between polymer structure or water content and gel strength but gels based on vinylpyrrolidone do tend to be stronger. Work on cross-linked acrylic acid gels for microfluidics showed similar strengths, with a dramatic decrease as the gel was swollen at high pH [37]. A cross-linked copolymer gel of hydrophilic and hydrophobic segments was reported to have a strength of 200–500 kPa [38]. 1.2.5

Multi-Phase Gels

Many gels are two-phase composite systems. Polyacrylamide gels are often quite turbid, suggesting phase separation into polymer-rich and polymer-poor regions. The cross-linked structure prevents large-scale separation, so unambiguous evidence for two phases is hard to obtain. Crystallizable synthetic polymers form solvent-containing gels, which apparently contain crystallites connected by segments of solubilized polymer chain. Similar combinations of regions of nanoscale order linked by disordered solution probably characterize many biological gels, such as gelatin, agarose and calcium alginate. In principle the phase behavior of a lightly cross-linked gel would be expected to be the same as that for a high molecular weight sample of the same polymer in the same solvent. Heavier cross-linking would restrict the entropy of the chain and might induce phase separation. The search for an artificial cartilage material has long driven the search for strong gels. Various multi-phase systems have been found that are much better than simple gels but, until recent unexpected results on ‘double network’ gels, none have been strong enough to look really promising. It is has been known for some time that the properties of poly(vinyl alcohol) and mixed polyacrylic acid/poly(vinyl alcohol) gels can be enhanced by a series of freeze–thaw cycles that drive more extensive aggregation of the polymer [39]. Exactly what happens is unclear but growth of ice crystals will probably concentrate the polymer to the crystal boundaries and drive formation of insoluble hydrogen-bonded complexes of the polymers [40]. Addition of DMSO as a co-solvent enhances the gel strength, possibly by limiting ice crystal size. Early work on two-phase freeze–thaw modified neutral gels of poly(vinyl alcohol) mixed with cationic and anionic polymers found a strength of 1 MPa at 85 % water [41]. More recently, such poly(vinyl alcohol) gels with water contents of around 80 % were found to fail in compression at a few MPa [42]. This freeze–thaw process produces a two-phase composite structure which has recently been studied in more detail [43, 44]. Composite gels with poly(vinyl alcohol) and other water-soluble polymers have also been studied [45].

Polymer Gel Actuators: Fundamentals

13

There have been many recent studies of composite gels made by irradiation of mixed solutions of polymers and increases in strength have been reported when compared to single polymer gels [46]. It would be expected that the properties of these disordered systems would primarily follow the water content. A number of studies have considered reinforcement of gels with inorganic fibers or plates both added before gelation and grown in situ in the gel and a significant increase in modulus is certainly seen [47–49]. With exfoliated clays as reinforcement, moduli increased from 4 to 20 kPa as the clay was added and the tensile strength increased from 0.1 MPa to 0.3 MPa [50–52]. One very attractive approach, based on the analogy to collagen-reinforced biological gels, is to reinforce gels with fibrils of rigid-rod polymers [53]. This particular system does show a significant increase in modulus but from very low values and no strength data was given. Thus the full potential of composites of this type has yet to be fully explored. A simple variant of this approach is to reinforce a gel with a textile, such as non-woven polypropylene [54, 55]. 1.2.6

Double Network Gels

Gong et al. [56] have formed gels with compressive strengths up to 17 MPa at 90 % water content by forming an interpenetrating network of ionic and nonionic gels in a two-step process (Figure 1.1). This compares with a strength of 0.2 MPa in compression for the equivalent single-component gel. These gels are produced by forming a moderately tightly cross-linked network, then swelling this gel in a solution of a second monomer with a low ratio of cross-linking agent and carrying out a second polymerization. As a result of the high degree of swelling in the monomer solution, the first gel network is highly extended in the final product while the second network is relaxed. The weight fraction of the second network in the final gel is 10–20 times that of the first network. 20

PAMPS-PAAm DN gel

Stress (MPa)

15

10 PAAm gel PAMPS gel

5

0 0

20

40

60

80

100

Strain (%)

Figure 1.1 Stress–strain curves from DN gels show much higher strength than conventional gels (Reprinted with permission from Gong, J. P., Katsuyama, Y., Kurokawa, T. and Osada, Y. Double network hydrogels with extremely high mechanical strength, Advanced Materials, 15, 1155–58, Copyright (2003) Wiley-VCH Verlag GmbH).

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Biomedical Applications of Electroactive Polymer Actuators

Other gels with a more lightly cross-linked first network show yield and necking in tension with extensions to break over 10  (1000 % þ strain) and a strength of 0.3 MPa [57]. Using a tear test, a fracture energy of 300 J/m2 was measured, compared to 0.1–1 J/m2 for conventional gels [58]. Yield is characteristic of metals and semicrystalline polymers where molecular slipping sets in at high stresses. It is not normally seen in rubbers or gels where the permanent cross-linked network prevents slippage. While strength was greatly increased by the addition of the second component, the initial modulus was affected little when compared to a conventional gel at the same concentration. If these double gels are made with a linear polymer in place of the second network, the fracture strength and fracture energy rise steeply at a molecular weight over 106, where the second polymer becomes highly entangled and these entanglements can act as physical cross-links [59]. Cyclic loading tests do show hysteresis, with a loss of modulus after successive cycles to high strain. This implies that breakage or other irreversible loss of some highly strained cross-links is occurring [60]. Other workers have found similar enhanced strengths in poly(ethylene oxide)–polyacrylic acid double network gels [61, 62]. These gels have tensile strengths which range up to 12 MPa depending on composition and swelling. Molecular dynamics simulations have been carried out for these PEO–PAA gels and show that the elastic modulus rises suddenly at strains of about 100 % where the first network becomes fully stretched [63]. This combination of high stress with a large strain to break increases the fracture toughness. Other interpenetrating networks, formed without the extension of the first network, have not shown similar improvements in strength. For instance HEMA–gelatin gels reach strengths of 65 kPa, just slightly above gelatin alone [64] and polyacrylamide/polyN-isopropylacrylamide IPN gels reach strengths of just 10 kPa [65]. Templating gels on a colloidal crystal and then removing the colloid has also been reported to give good toughness, although the modulus and strength remain low [66]. Reinforcing polyacrylamide gels with a rigid-chain polyelectrolyte does lead to a large increase in modulus [53]. Shull has recently emphasized that there is a need to develop better understanding of fracture toughness in these double network gels and in biological gels [67]. The preceding discussion also shows that microstructure control can lead to greatly enhanced strength and toughness in gels. Improvements obtained by double networks, by freeze–thaw and by fiber reinforcement suggest that there are many possible routes to better properties. A J-shaped stress–strain curve, that is an increase in elastic modulus at high strain, seems to be one signature of better toughness and strength. In this view, cartilage is a similar cross-linked network of collagen microfibers with a second network of coiled proteoglycan chains that can absorb fracture energy. Thus it seems that modulus and strength of networks can be separately controlled in order to design suitable mechanical performance into any functional matrix [68]. A useful objective for future work would be to develop some design rules. 1.2.7

Transport Properties

Many potential applications of active gels, as muscles, for drug release or as sensors, depend on their ability to respond to external influences by changing volume or shape or by taking up or releasing small molecules. If this responsiveness is not important, their mechanical properties can be duplicated by a range of dense elastomers and there is no

Polymer Gel Actuators: Fundamentals

15

reason to employ a gel. Small molecule transport properties are thus crucial. Since drug release has been well discussed earlier [69], the focus here is on the faster diffusion times appropriate for sensing and actuation. Iit would be expected that diffusion coefficients of small solutes in gels should be intermediate between those in solution and in an elastomer. The diffusion coefficients of solutes in dilute gels have been measured and do not differ dramatically from those in solution [70]. Diffusion processes in gels can also be conveniently studied by conductivity [42]. For low levels of soluble small molecule additives in an elastomer, the diffusion coefficient can be estimated from the properties of the polymer and the size of the molecule [71]. At the other extreme, of a highly swollen hydrogel, the diffusion of water and of soluble compounds in the water have been shown to be reduced roughly in proportion to the water content of the gel. A significant difference will result when the diffusion process causes swelling or deswelling of the gel. The resulting nonuniform volume changes through the gel will result in highly non-Fickian behavior and may also cause significant internal stresses or fracture. Tanaka and co-workers [72] studied many of these interactions. As gel concentration changes in a single-phase gel, diffusion and solubility of a solute will change, such that the resulting changes in permeability can be quite complex. Two-phase gels will be even more complicated and so should be a source of many complex changes in release or uptake of solutes. In building a gel based machine, it will probably be equally desirable that the gel can swell and deswell over a range without dramatic changes in the properties of the surface region. This requirement imposes a coupling of effective sample size, diffusion rate and response time on any gel device. The self-diffusion coefficient of water is 2  10–5 cm2s–1 and diffusion times can be roughly estimated as x2/2D, where x is the effective half-thickness of the gel. This gives a response time for a gel one millimeter thick of about one minute, which would probably be the longest acceptable response time for many devices. It is clear that any responsive structure that is larger could have a finely divided spongelike structure that allows for fluids to flow in and out through channels rather than simply diffusing. If a hypothetical microstructural scale of 100 microns thickness is adopted, a diffusion coefficient of about 3  10–7 cm2s 1 can be accepted in a working device. 1.2.8

Drying

Engineering materials are generally dry and there is a reluctance to use materials that may lose liquid and dry out. However, we do work with systems that need to retain liquids, such as foods, cosmetics, paints and inks. Our experience with houseplants suggests that it would be quite possible to develop long-lived systems that depend on occasional replenishment of a water reservoir if there were desirable and unique properties. This survey of recent work on gels focuses on whether it is possible to reach a combination of mechanical properties, stability and activity that would allow more use of gel devices and structures. Drying time is also clearly an issue in gel devices. The evaporation rate of water is very dependent on temperature, humidity and air flow but measurements on drying of snails give a rate of about 100 microns/h as typical for still air with an active snail

16

Biomedical Applications of Electroactive Polymer Actuators

being able to reduce this by about 20-fold by maintaining a surface coating [73]. If some similar mechanism was available to gel devices with a size of about one centimeter, they would experience 10 % water loss in a week and so would need only occasional rehydration.

1.3 Chemical and Physical Formation of Gels While linear polymers can be purchased in bulk and processed in the melt or solution to the desired form, cross-linked gels must be chemically formed in situ. Gel actuators will have to be built as systems containing multiple layers or material, structured pores and electrodes. Processing methods that lend themselves to building these systems must be available. Many gels, are chemically cross-linked like vulcanized rubbers and this limits our ability to process shape them. Just as thermoplastic elastomers have allowed an expansion of the range of applications of rubbery materials in complex shapes, so meltable gels, like agarose, can be more readily shaped than chemically cross-linked gels. Most of the work on synthetic gels uses gels formed by free radical polymerization of the families of hydrophilic acrylate, methacrylate and acrylamide monomers plus vinylpyrrolidone. While these methods give a very versatile family of hydrogels, it is worth noting that the polymers are atactic or otherwise irregular and this limits formation of any microstructure that might give rise to toughness. Also, free radical polymerization is oxygen sensitive and this can make it difficult to get good control of the polymerization in small or air-exposed devices. For these reasons, it may be valuable to explore other approaches to forming synthetic hydrogels. The kinetics of free radical linear polymerizations has been thoroughly studied [74] and the relationships between molecular weight distribution and polymerization conditions are well known. Gels are made by incorporating a small fraction of bi-functional or multifunctional monomers that becomes part of more than one kinetic chain so that a network forms. The statistics of network formation are also well known. Two factors act against being able to make reproducible samples of muscles. These reactions are very sensitive to small amounts of impurities as well as oxygen, and these can be difficult to control in small samples. In addition, the structure of a network is much more difficult to characterize in detail than is a linear polymer. As a result, it is hard to know what is the real structure of a gel sample and usually they are not characterized to any significant extent. Many gel samples are formed by UV irradiation. Quite reproducible samples can be formed by irradiation of de-oxygenated samples held between glass plates under conditions where the UV is only slightly attenuated as it passes through the sample (Figure 1.2). On the other hand, samples made in air under strong UV will tend to have structures that change properties through the thickness and many residual unpolymerized chain ends in the network. The reaction kinetics will depend on the water content of the monomer solution during polymerization. This, in turn, will affect the network structure and so the swellability of the final gel network. Many studies of gel actuators use gels with low cross-link densities that show very large equilibrium swelling. To get good mechanical properties, it is preferable to work with gels that are more like contact lenses and only swell to become about 50 % water at equilibrium.

Polymer Gel Actuators: Fundamentals

17

Figure 1.2 Gel microcantilever in water before (a) and after (b) UV irradiation (Reprinted with permission from Watanabe, T., Akiyama, M., Totani, K. et al. Photoresponsive hydrogel microstructure fabricated by two-photon initiated polymerization, Advanced Functional Materials, 12, 9, 611–4, Copyright (2002) Wiley-VCH Verlag GmbH).

To avoid the oxygen-sensitivity problem, Yoshioka and Calvert studied epoxy hydrogels for small artificial muscles and sensors (Figure 1.3) [75, 76]. Water-soluble polymers and hydrogels have also been made by ring-opening metathesis polymerization [77–80]. It would also be expected that it is possible to form stable hydrogels based on polyamides and other polymers formed by condensation chemistry that might form tough microstructures. 1.8

Thickness Changes (h/ho)

1.7 Gel 4

1.6 1.5 1.4 1.3 Gel 3

1.2 1.1 1 0

120

240

360

480

600 720 840 Time (s)

960 1080 1200 1320

Figure 1.3 Swelling–collapse behavior of two epoxy gels in 0.01 M NaCl at pH 4.3; gel dots are printed on the platinum plate, which contacted the anode side. Average dot diameter (d0) was approximately 460 mm and the initial thickness (h0) was 130 mm. Applied voltage was 6.0 V and current was 3.6 mA (Reprinted from dissertation work of Y. Yoshioka, Courtesy of University of Arizona).

18

Biomedical Applications of Electroactive Polymer Actuators

It has long been know that many pairs of polymers self-assemble to form gels through interchain bonding. Gelled capsules can be formed by dripping a solution of cationic polymer into a solution of anionic polymer [81, 82] or by dripping alginate into calcium [83]. This process results in a thick-walled capsule with a nonuniform structure. There is no way of simply mixing the two components so that a uniform block of material is formed. Ionic selfassembly by sequential dipping into anionic and cationic polymers [84] or by repeated contact printing [85] does produce uniform thin films of ionic gels in a more controlled fashion. While these systems have been shown to have many potential applications, there has been little work on the structure and properties of the gels themselves. One structural study is from Lewis et al. [86] where a fine capillary was used to make micron-scale 3D structures by extruding a stream of solution into water at a pH that induced gelation (Figure 1.4). Such writing systems could be used to make a wide range of gel microdevices.

Figure 1.4 Printed polyelectrolyte complex filaments, written with a micron-sized nozzle. Scale bar ¼ 10 mm (Reprinted with permission from Gratson, G. M., Xu, M and Lewis, J. A. Microperiodic structures: Direct writing of three-dimensional webs, Nature, 428, 386, Copyright (2004) Nature Publishing Group).

Natural proteins form structures by a combination of ionic interactions, binding to multivalent cations and hydrogen bonding. It is possible to design synthetic proteins to form similar structures [87–89] and demonstrate that they gel. Thus there would seem to be considerable opportunities for better characterization of gelation of synthetic ionic

Polymer Gel Actuators: Fundamentals

19

polymers and for the study of more structured synthetic gel systems formed by other polymerization chemistries. This should lead to a better understanding of the whole structure–mechanical properties map for gels.

1.4 Actuation Methods The merit of a muscle-like actuator depends on a fast response time, a high contractile force and a large actuation strain. The state of swelling of a gel is dependent on temperature, solvation and pH. These factors may in turn be influenced by the application of chemical changes, light, heat, electrical and magnetic fields. The most desirable way to control a gel actuator is via an electric field but electrically driven actuators have so far produced little force. The simplest form of control, for understanding the response, is to change the temperature. 1.4.1

Thermally Driven Gel Actuators

Polymer chain coils in solution will undergo conformations changes as the temperature changes [90]. In many cases, the coil contracts as the temperature is raised until the solution separates into polymer-rich and dilute phases at the lower critical solution temperature (LCST). A homopolymer gel with a low density of cross-links is thermodynamically similar to a solution of very high molecular weight polymer. Thus, the gel will contract under conditions where the linear polymer coils contract. Tanaka studied this process in gels using the concepts of phase transitions and distinguishing between gels which show a continuous change from the expanded to the contracted state and those which go through a phase separation [91]. Poly-N-isopropylacrylamide (NIPAM) solution in water precipitates above 40 °C, so NIPAM gels have provided the most fruitful example of a thermally driven phase change. A swollen cross-linked NIPAM gel deswells at the same temperature, 40 °C. Hompolymer gels show a discontinuous phase transition while copolymer gels with ionizable groups such as acrylic acid can show a continuous transition or critical point behavior. Having resolved the thermodynamics of the system, the kinetics of the phase change become important. The kinetics of gel swelling can be treated as a two-step process. Solvent diffuses into the gel causing some regions to swell and then there is an instantaneous shape change to minimize the elastic energy between the swollen and unswollen regions. For swelling involving small changes in volume in the continuous region of the phase diagram, the kinetics and dependence on the gel shape correspond to those expected for diffusion of solvent. Thus the response time of a long acrylamide gel cylinder 1.3 mm in diameter was about one hour, in agreement with a theory based on the coupling of the diffusion coefficient of water in the gel and the shear modulus of the gel [92]. In the two-phase region, or where there is a large volume change, the changes in diffusion coefficient and gel properties with the local state of swelling will lead to complex kinetics which depend on the details of the initial and final gel states. A number of workers have addressed the diffusion time by making the gels porous, so that simple diffusion only occurs over a small distance. The effective diffusion rate can be

20

Biomedical Applications of Electroactive Polymer Actuators

linked to the degree of porosity [93]. For instance, polymerizing the gels under reduced pressure produces a macroporous gel that responds to temperature changes in a few minutes, about 10 times faster than normal gels [94]. After freeze drying this porous gel had an apparent pore size of 20 microns although the actual structure prior to drying was not studied. Viewed as an actuator, these porous systems have the disadvantage that the forces developed will also be reduced as the porosity increases. Polymerizing gels in the presence of poly(ethylene oxide) also yields porous, fast-responding gels [95], as does polymerization of gels on micron-scale liquid templates [96]. A related, second approach to increasing response time is to prepare a two-phase gel so that a nonresponsive matrix can allow fluid flow into and out from the gel. This has been shown for solutions of linear NIPAM as a block or graft copolymer with poly(ethylene oxide) [97]. The collapsed, precipitated state of the graft copolymer is more open than pure NIPAM and so allows more rapid water penetration and redissolution when the temperature is decreased. This type of approach addresses the problem that a rapidly shrinking gel tends to form a dense skin that inhibits water loss from the interior, both slowing the volume change and possibly leading to fracture of the gel under the resulting shrinkage stresses. The effect of such skin formation on deswelling kinetics has been studied by Hirose and Shibayama, who showed that pure NIPAM gels form a dense layer and shrink much more slowly than weakly charged copolymers of NIPAM and acrylic acid that retain more mobility for water in the collapsed state [98]. NIPAM has been most widely discussed but other aqueous gels are also thermally responsive, including poly N-vinylcaprolactam (PVCL) and hydroxypropylcellulose (HPC) [99]. In the unswollen (high temperature) state, PVCL is much stiffer than in the swollen state and resembles a glassy polymer, whereas in HPC the modulus decreases as it shrinks. These thermally driven gels have been widely studied with a view to use as actuators or in drug delivery. While the volume change is large, the response is slow, primarily because water must diffuse into and out from the gel. In addition, rapid heating may be quite easy to achieve but it can be difficult to cool the gel rapidly in any practical fashion. The stress developed on shrinkage can be estimated from a knowledge of the relationship between elastic modulus and degree of swelling, as discussed below. 1.4.2

Chemically Driven Gel Actuators

While electrical actuation would be the most practical method to control and drive gel actuators, gels primarily respond to the chemistry of their environment and so chemical actuation has proved to be the most efficient method to drive gel actuators. Muscle is a chemically driven actuator where the energy of the ATP to ADP conversion is used to drive a cyclic shape change in myosin that causes it to ‘walk’ along actin filaments. From the viewpoint of building an artificial muscle, it is important to note the points that: i. the energy to drive this process in the short term is stored locally in the cells as glucose and oxygen; ii. muscle develops force only while it is burning energy and will passively extend once the cells cease to produce ATP. It does not move between two states as gels do. Thus the analogies between real muscle and chemically contracting gels are not as close as they might seem [3, 4, 100].

Polymer Gel Actuators: Fundamentals

21

In most cases the source of chemical energy for gel actuators is either a change in pH or a change in solvent. Many gels are based on cross-linked polyacrylic acid. The linear polymer is very soluble in water at pH values above about five where it is in the ionized form as a sodium or other salt. At low pH, below four, the acid form predominates and is only slightly soluble in water. Repulsion effects between adjacent ions on the polymer chain cause the pH range over which the polymer goes from the acid to salt form to be much wider than for a simple acid such as acetic acid [90]. The chain is also much more extended due to Coulombic repulsions in the salt form. The chain extension is sensitive to the ionic strength of the aqueous solution, since other ions in solution will screen the repulsions and allow the chain to coil. In terms of the phase transition treatment of Tanaka discussed above [72, 101], both commonly used gel muscles, polyacrylic acid and polymethacrylic acid, are soluble in water at all pH values and so the acid to ionized change is a continuous phase transition in water. In mixed solvents the change can be discontinuous. The response of polyacrylic acid and polymethacrylic acid gels to changes in pH and solvent have been widely studied [102]. Other acid gels [103] and chitosan gels [104] have also been studied. However, the number of systems explored only represents a tiny sample of the potential range of chemically actuating gels. These two monomers can be easily copolymerized with a wide range of other monomers to form gels with different solubility characteristics. It can be expected that: addition of a more hydrophobic monomer will induce phase separation at low pH; addition of a strongly ionized monomer, such as a sulfonated monomer, will limit shrinkage at low pH; and addition of a neutral water soluble monomer, such as hydroxyethylmethacrylate, will limit the maximum swelling at high pH. These remarks apply to random copolymers. Graft or block copolymers could be made with a wide range of responses to pH, ionic strength or solvent. An amine-functional monomer such as aminoethylmethacrylate can be expected to respond as a mirror-image to the acidic gels, swelling at low pH and expanding at high pH. Yoshioka and Calvert have studied the response of epoxy–amino gels [75, 105]. Amine gels could also be formed with a wide range of comonomers. One interesting but littlestudied case is amphoteric gels containing both amino and acid groups which may be swollen in acid, in base or in salt solutions [106–108]. Other chemistries may also be used to reversibly change the properties of hydrogels. Thus, gels with reversible disulfide cross-links have been studied [109–111]. Reversible metal-ligand cross-links with a self-repairing function have also been described [112]. Suitable modification of these systems could also be used to form chemically driven actuators. In most studies, pH is adjusted by exchange with a mineral acid such as hydrochloric acid and sodium hydroxide. Ammonium and potassium counter-ions would be expected to act very similarly but lithium or divalent ions, such as calcium, will not result in such large solubility changes. Lightly cross-linked gels with more than about 20 monomer units between cross-links will be expected to act as a polymer of infinite molecular weight as it responds to the medium by expanding and contracting. Higher cross-link densities will decrease the tendency of the gel to swell, both by limiting the extent to which the chains can coil and uncoil but also by limiting the solubility of the chains. Gels with more acidic side groups, such as sulfonated gels, will be less responsive because very low pH would be needed to deionize the gel. Phosphated gels would be expected to show several ionization steps but have not been much studied. In addition to

22

Biomedical Applications of Electroactive Polymer Actuators

acid gels, gels with amine side groups respond to pH. They show a mirror of the behavior of acids, contracting in strong base but ionizing and expanding in acid [75]. Gels will also expand and contract as a result of being moved between water and a watermiscible organic solvent such as acetone. The expansion and contraction of the gel reflects the solubilization and precipitation of the equivalent linear polymer in the solvents. The best performance of chemically driven gels is by polyacrylonitrile fibers which are partially converted by heat and alkali to cross-linked polyacrylic acid fibers. Previous work has shown that these can be made into muscle-like actuators that respond to pH changes with good force and strain characteristics [6, 113]. The good characteristics of this material arise because the spun fibers are oriented and so the swollen gels are stronger than unoriented gels. Orientation must also affect the actuation properties but this does not seem to have been studied. Also, the fibers have a small diameter and so respond rapidly to pH changes. Recent work has found that commercial poly(acrylonitrile) (PAN) yarns treated in this way could give actuator strains of about 80 % and stresses of 100 kPa to 1 MPa with a response time of about 10 seconds [114]. The same fibers can be driven electrically at 5 V, using an electrode embedded in or adjacent to the fiber, to produce similar stress but the response time is much longer, about 10 minutes [113, 114]. Electrospinning has been used to make fibers of less than one micron diameter that could then be twisted into yarns [115]. Actuator strains of 38 % were obtained in cycling from pH 1 to pH 12 with a response time of about five seconds. The authors claim an actuator stress of about 10 kPa but the data presented suggest this value is actually closer to 1 MPa. Other similar stiff synthetic gels have also developed higher forces in response to chemical activation [116]. Actuation that relies on pH changes has the advantage that hydrogen ions diffuse rapidly. In biology, changes in calcium ion concentration often drive changes in molecular conformation. Forisomes, plant proteins responsible for opening and closing of leaf pores, produce a force of 11 kPa in response to calcium and pH [117]. An actuator based on changes in copper ion concentration has been described [118]. As the copper was oxidized to Cu2+ and reduced back to copper metal, the gel contracted and swelled. The oxidation and reduction could be driven chemically or electrochemically. 1.4.3

Gels Driven by Oscillating Reactions

Several groups have recently studied self-oscillating gels that swell and contract cyclically in response to an oscillating chemical reaction in the medium. Yoshida and co-workers have demonstrated patterns of contraction in NIPAM gels that cycle through the lower critical solution temperature (LCST) as the gel cycles between two oxidation states [119]. Yashin and Balasz have modeled the moving patterns of expansion and contraction set up in these gels [120]. Jones and co-workers have used oscillating pH reactions in the liquid surrounding an acidic block copolymer to drive a gel motor [121, 122]. However, their calculations of power density for these gels do give very low values. This approach does suggest that it might be possible to use an electrical signal to set off such a chemically driven oscillation. This would be a good mimic of natural muscle, where an electrical impulse triggers chemically driven contraction of the muscle.

Polymer Gel Actuators: Fundamentals

1.4.4

23

Light Actuated Gels

With the availability of powerful LEDs and solid state lasers, light would seem to be an excellent way of communicating with gels in order to drive actuators or read out from sensors. Light can be used to heat an absorbing gel and cause a shape or volume change. Light could also be used to change pH in order to produce a volume change in a gel. Gels can also be made with light-sensitive groups that undergo a chemical conversion that results in a shape change. So far none of these systems has shown a strong enough response to be attractive as an actuator. Light can also be used as a heat source to swell and deswell thermally-sensitive gels. Suzuki and co-workers showed that a NIPAM gel containing copper chlorophyllin, a soluble dye derived from chlorophyll, can be induced to swell and contact as the light is turned on and off [123]. A copolymer of NIPAM with sodium polyacrylate showed similar behavior but with a large hysteresis, such that the gel could be switched to the unswollen state at high intensity but would remain in either the swollen or unswollen state over an intermediate intensity range [124]. Similar bi-stable behavior was shown with pH or temperature. In principle, this type of bi-stable switching would be of interest for many applications. The bi-stability was analyzed in terms of a Landau model for phase transitions and might also be viewed as relating to skin effects. In lieu of a dissolved dye, a similar response to heating from a near-IR laser was obtained in NIPAM gels containing dispersed gold nanorods [125]. Beebe and co-workers have produced gels containing gold nanoparticles that respond by swelling on absorption of selected wavelengths of light and so act as valves to open or close channels in a microfluidics system [126]. Similar gold-hydrogel composites also show a change to becoming electrically conducting as the collapse [127]. The photoisomerization of azobenzene between the cis and trans forms is well known and gels containing attached azobenzene units have been shown to respond to UV illumination by stiffening [128]. Leucocyanides are photoresponsive dyes that convert between ionized and unionized forms on irradiation with light and cause osmotic swelling and contraction of a gel [129]. Marder and co-workers developed a hydrogel that responds to UV irradiation with a keto to enol tautomerization that results in mechanical deflection of a cantilever [130]. Nitrocinnamate chemistry has been used to create a hydrogel which can be reversibly photocross-linked and photocleaved [131]. A copper cross-linked polyacrylic acid gel, containing titanium dioxide particles, has been shown to swell under UV light and contract again in the dark [132]. A similar system with silver-coated titanium dioxide has also been described [133]. The volume change is large but the response time is hundreds of minutes. 1.4.5

Electrically Driven Gel Actuators

A recent review [134] of electroactive gels focused particularly on polyelectrolyte gels based on cross-linked polymers of acrylic acid. Such acidic hydrogels can show a large volume change in water as the pH changes from acidic to basic and the gel becomes ionized. A pair of electrodes in a suitable salt solution will produce hydrogen at the cathode and oxygen at the anode. This results in decreasing the pH at the cathode and increasing pH at the anode. A pH-sensitive acid gel will thus shrink if it is attached to the cathode in a

24

Biomedical Applications of Electroactive Polymer Actuators

75

0.2

50

0.1

25

Strain

0.3

0

5 Time (min)

10

Weight gain (%)

solution at neutral pH. Depending on whether one or both electrodes are embedded in the gel or are in the surrounding solution, a gel may shrink, expand or bend [135, 136]. Several groups have demonstrated actuation by these gels, while immersed in solution, with embedded or external electrodes. Osada and co-workers made gel ‘loopers’ that would crawl along a bar between electrodes [137, 138]. Shiga and co-workers carried out an extensive series of studies on poly(vinyl alcohol)– polyacrylic acid gel actuators driven electrically (Figure 1.5) [136, 139–143]. These gels are prepared by a freeze–thaw method that produces a strong and highly porous structure. A ‘fish’ was made that would swim along a channel between external electrodes (Figure 1.6) [39]. Although the forces are small, such gels can be used to transport ‘cargo’ as they expand and contract in a tube [144].

0

Figure 1.5 Strain in bending and weight gain of poly(vinyl alcohol)–poly(sodium acrylate) composite hydrogel under an electric field of 10 Vcm–1 (Reprinted with permission from Shiga, T., Hirose, Y., Okada, A. and Kurauchi, T. Bending of poly(vinyl alcohol)-poly(sodium acrylate) composite hydrogel in electric fields Journal of Applied Polymer Science, 44, 249–53, Copyright (1992) Wiley-VCH Verlag GmbH).

Natural muscle undergoes a linear contraction. A single-component gel placed symmetrically between electrodes will bend as one side expands and one contracts. A combination of layers of two different gels that expands and contracts in a linear fashion when the cathode is embedded in one layer and the anode in the other has been demonstrated [145]. This can be achieved by having gels of opposite ionic charge or by having a high modulus contractile gel layered onto a soft neutral gel. In this second case, as the stiff gel contracts the water is pushed into the neutral, soft gel and this expand in thickness. Almost all gel actuators work in solution or with embedded electrodes but gels have been demonstrated which respond to large electrical fields in air [146]. The gels showed a crawling motion driven by electrostatic response of the gels to the applied field.

Polymer Gel Actuators: Fundamentals

25

Figure 1.6 An ‘eel’ gel of PAMPS moving in an oscillating electric field (Reprinted with permission from Osida, Y., Okuzaki, H. and Hori, H. A polymer gel with electrically driven motility, Nature, 355, 6357, 242–4, Copyright (1992) Nature Publishing Group).

1.4.6

Electro- and Magneto-Rheological Composites

Recently there has been much interest in composites of magnetic particles in a soft matrix which respond to a magnetic field by a change of shape and properties [147]. These materials are the elastomeric equivalents of magneto-rheological fluids and electrorheological fluids. The prototypical system is a dispersion of 25 % of micron-sized iron particles in gel of silicone rubber and silicone oil but polyurethanes and other rubbers have been investigated. When cured in a magnetic field the particles form chains [148]. After curing, the stiffness of the material increases in a magnetic field. At a strain of 5 %, an increase in stress of about 50 % is seen in a field of 123 kA/m (1500 Oe). In a study of magnetostriction, a field of 800 kA/m (1 Tesla) produced a strain of 0.3 % in a sample preloaded in compression to 100 kPa, a strain comparable to magnetostrictive alloys (Terfenol-D) [149]. These materials also show an electric field response if they can be formed to be nonconductive [150]. A shear stress of a few kPa can be produced at a field of 1 MV/m. A study on barium ferrite particles in carrageenan hydrogels showed a modulus decrease of about 75 % (from 20 MPa) on magnetization at 800 KA/m [151]. Strains of about 0.2 % could be produced in these gels. This effect is clearly not the same as that seen in the ironfilled elastomers since it is larger, in the opposite direction and irreversible.

26

1.4.7

Biomedical Applications of Electroactive Polymer Actuators

LC Elastomers

Liquid crystalline (LC) elastomers show a contraction on heating from the nematic to isotropic phase and so can be pictured as actuators [152, 153]. Similar behavior can also be seen in liquid crystalline gels [154], which opens the possibility for combining electrical switchability, the shape change associated with liquid crystalline elastomers with the volume change associated with gels [155, 156]. Similar changes have been demonstrated in hydrogels [157]. These systems can deliver rapid large shape changes, stresses of 130 kPa have been reported for thermal actuation [158], which is enough to be useful as actuators.

1.5 Performance of Gels as Actuators The performance of muscle has been measured well on scallop [159]. Muscle is characterized by a response time of less than one second. The maximum actuator stress is about 300 kPa at a mid-point of contraction, corresponding to about 12 % shortening. The maximum strain is 25 %. The energy density of muscle is about 50 kJ.m–3. The power output is 200 W/kg peak or 50 W/kg for sustained cyclical contraction. Reaching or exceeding these characteristics is now the goal for an artificial muscle system. Unlike many mechanical systems, muscle uses power as long as stress is needed to hold a static position. As a result, the normal mechanical concept of efficiency is hard to apply. For instance, the position of the arm may be held by balanced force between a muscle pair trying to extend and contract the elbow. No motion occurs, but considerable energy is expended in this isometric exercise. As a result, discussions of efficiency can be misleading, since the conditions of measurement need to be well defined. The actin–myosin system is essentially the only type of actuator in animals but there are a few other examples of muscle-like tissues. The jelly bodies of jellyfish, sea anenomes and other sea animals contain proteoglycan hydrogel, elastic fibers, collagen and muscle fibers in various arrangements [160–162]. The sub-micron reinforcing fibers are loosely connected to the matrix and so provide strength at high extensions and slow the viscoelastic response of the tissue [8]. In sea cucumbers, starfish and other echinoderms, the mechanical properties of the gel can be quickly changed from soft to hard. This was thought to occur by release of calcium ions to harden the polyanionic matrix but actually seems to be due to the release of proteins that temporarily bond to and cross-link the collagen fibrillar network [163]. This response is essentially a change in elastic modulus, as opposed to an active contraction such as is characteristic of muscle. In legumes, specialized cells open and close to control fluid flow in the vascular system of the plant. These ‘forisomes’, which are 30 microns in diameter, contract in response to calcium ions in about 50 milliseconds and develop a force of 11 kPa [117]. The available energy density has been estimated at 0.5 MJ.m–3, which is close to what would be needed for an artificial muscle [164], and preliminary tests have been made of forisomes in a microfluidic system [165]. It may be that this system will provide a better model for artificial muscle than the much more complex actin–myosin system. The target performance characteristics for a biomimetic muscle would be a response time of about one second, an actuator strain of about 25 % and a developed stress of about

Polymer Gel Actuators: Fundamentals

27

300 kPa. In the absence of any current good answers, any approximation to these values would be of interest. At present, the only systems that approach this are the converted PAN fibers acting as chemical actuators. The actuation stress for a gel of known elastic modulus can readily be estimated. Consider, for example, a gel that initially contains about 50 % water. This is a much lower water content than most experimental hydrogels but is similar to a contact lens formulation and would have an elastic modulus in the region of 1 MPa [36]. Assume the gel will act as an actuator in compression, by pushing a weight upwards. A change in thickness of 25 % would correspond to a volume change of 2, assuming width, length and thickness change equally. To develop a substantial force at this new thickness, the volume change would have to be greater, for instance to 3.4, which corresponds to an unconstrained thickness change of 50 %. Using the equation for modulus and swelling above, the elastic modulus will have dropped to about 660 kPa at this swelling. When this swollen gel deforms elastically back to a thickness of 25 % greater than the original value, the stress generated will be 165 kPa. In contrast, more lightly cross-linked gels, with water contents of 90 % or more, have moduli of a few kPa and can exert little force in this format [166]. This then shows that an elastic gel, acting as an actuator in tension or compression, can only generate a substantial force if the elastic modulus is quite high and the water content relatively low. The swelling pressure exerted by these dilute gels can be much higher than the modulus [20, 166]. This can be up to hundreds of kPa at low degrees of swelling but drops rapidly as the polymer swells toward equilibrium. Ionic gels show higher swelling pressures under conditions where they are charged [21]. To use this to drive an actuator, the gel would need to be confined, for instance in a rigid cylinder with porous walls and a moveable piston. It should also be possible to exploit the swelling pressure in a suitable composite structure or oriented structure where swelling is free along one axis but is very constrained in the other two directions. Some of the marine animal structures, with reinforcing fibers wound spirally at an angle to one axis, may fulfill this requirement by only allowing expansion along one axis. These preceding arguments apply however the change in gel swelling is brought about, whether through thermal, chemical or electrical energy input. Thermal activation provides a convenient method for driving the actuator, which can clearly be engineered by a number of different routes. Cooling will be very slow for actuation occurring at 40 °C but a higher temperature could allow both rapid heating and cooling. Nonetheless, the associated systems for removing heat could add substantially to the weight and volume. It is also not clear that the phase changes necessary for thermal actuation can occur in gels at the high solids loadings needed to produce a significant stress. Hinkley et al. report an estimated maximum work achievable by the thermally driven PVCL gels of 1 MJ.m–3 of dry gel, considerably higher than that for poly(vinyl alcohol) gels and comparable with muscle [99]. However, the response is very slow. This maximum work reflects the higher mechanical strength of these gels. In many other gels, the mechanical work done by thermal actuation is small because the gels have a low elastic modulus and low strength. A measure of the efficiency of these gels as actuators or transducers is the mechanical energy density and the power density, the rate at which energy can be delivered. Gels driven by solvent-induced contraction have achieved 135 J/kg but only 2 W/kg, which compares with 70 J/kg and 100–200 W/kg for muscle that is in turn similar to the energy storage density of a lithium battery (Figure 1.7) [167].

Biomedical Applications of Electroactive Polymer Actuators

wv [Nmm/cm3]

28

160

150/15

140

125/40

120

130/15

100 80 60 40 20 0 0

100

200

300

400

500

σ [kPa] Figure 1.7 Working energy of poly(vinyl alcohol)/poly(acrylic acid) hydrogels, cross-linked under different conditions of temperature/time, versus applied stress (Reprinted with permission from Arndt, K., Richter, A., Ludwig, S. et al. Poly(vinyl alcohol)/poly(acrylic acid) hydrogels: FT-IR spectroscopic characterization of crosslinking reaction and work at transition point, Acta Polymerica, 50, 383–90, Copyright (1999) Wiley-VCH Verlag GmbH).

Chemical actuation, as discussed above, can give a response with acceptable force, strain and speed parameters but has the drawback that delivering the fluid reagents requires inconvenient amounts of associated pumping and piping. In this context the human body can be viewed as a large amount of muscle with a roughly equal amount of piping to provide the driving energy and a supporting skeleton. Electrical actuation may be much more convenient in principle but brings a further set of engineering problems. The current flow necessary to ensure expansion or contraction of a gel can be estimated on the basis of the number of ions that must be inserted or removed to convert from the acid to the ionized state, of a carboxylate gel for instance. One gram of acrylic acid gel at 50 % by weight of water, contains 7 mM of acid groups, which corresponds to 670 Coulombs or 11 amperes for one minute. Any practical system will thus demand substantial currents, which in turn places requirements on the conductivity of the surrounding liquid. If the associated electrode reactions also produce gas, there will 75 cm3 of gas produced per gramme of gel switched. The previous estimate assumes that all the ions generated at the anode react with the surrounding gel. This is probably true if the electrodes are far apart, so that the acid generated at the anode and the base generated at the cathode do not diffuse but essentially remain in separate compartments. However, this wide separation will add to the volume and weight of the system and reduce the current flow at any given potential difference. A further design problem is that the electrodes themselves must be flexible enough to expand and contract without significantly restraining the volume changes of the gel. This can be achieved with flexible wire electrodes or some other type of soft electrode system. Thus, in the two-gel multi-layer system made by Liu and Calvert [145], the performance is still constrained by current flow and the ability of the gels to contract and expand over

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the wire electrodes. Likewise, the performance of electrically driven PAN fibers is quite inferior to that of similar fibers chemically driven [113] and it is hard to retain stability of the electrodes as the fibers cycle. The electrical response of many soft actuators has been demonstrated as a bending beam. Typically, a strip of gel will bend in alternating directions as the potential difference across the gel is cycled. These thin sheets typically show a fast response and large deflection. However, the actual strain in the surfaces of the gel can be quite small. It is less obvious how to assess these bending actuators as a muscle-mimic but one measure is the blocking force, the force delivered at the end of the actuator, as a fraction of the weight of the actuator. Ionic Polymer–Metal Composites (IPMCs), for example, can deliver a blocking force of about 50 the weight of the sample [168]. Muscle in linear contraction would deliver about 300 its own weight. The elastic modulus of most experimental gel actuators is so low that the blocking force also will be very low. Earlier work showed that the response of gels in an electric field depends on the composition of the gel itself, whether the gel is touching the electrode and on the composition of the solution [135]. More recently a number of finite element models have been developed to calculate the response of a gel to an electric field (Figure 1.8) [169–171]. These have the advantage of being able to take into account all the competing effects and provide a prediction for the response of a gel actuator under any conditions.

Figure 1.8 Bending curvature of hydrogels in an applied field, effect of gel modulus [200] (Reprinted with permission from Li, H., Ng, T. Y., Yew, Y. K. and Lam, K. Y. Meshless Modeling of pH-Sensitive Hydrogels Subjected to Coupled pH and Electric Field Stimuli: Young Modulus Effects and Case Studies, Macromol. Chem. Phys., 208, 1137–46, Copyright (2007) Wiley-VCH Verlag GmbH).

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It should be remembered that the chemical changes can both cause volume and stiffness changes, such that the results of combining pH changes and applied force can be unexpected [172]. This coupling between swelling thermodynamics and mechanical stress leads to a number of other peculiar phenomena, such as negative Poisson’s ratios [173–175], and may result in strange responses to complex loads, such as bending. Thus the challenge for gel actuators is to devise a practical solution to the whole system. Three possible approaches are:

 Develop a chemical actuator with a compact and reliable supply of the needed chemical energy, not necessarily depending on acid and base.

 Develop a thermal actuator based on expansion against a piston of a gel within a porous but rigid enclosure, similar to an automobile radiator thermostat.

 Develop a two-compartment fine scale electrically driven gel muscle as an extension of the electrically-driven PAN fibers.

1.6 Applications of Electroactive Gels Recent papers have explored producing faster responses by inducing porosity into the gels. While the application as artificial muscles is not practical, low stress applications, such as sensors and valves, are possible [176]. Lenses and light modulators have been demonstrated recently [177, 178]. 1.6.1

Gel Valves and Pumps

Active gels have been developed as valves for microfluidic systems, where swelling of the gel in response to light, pH, thermal or electrical stimulation can be used to close a fluid channel or act as a pump [37, 126, 176, 179–181]. Electrically driven gels essentially respond to local changes in pH or ionic strength caused by electrolysis of the water [179, 182]. Bassetti et al. show that, at high electrical fields, there is a fast response due to ion migration within the gel and then a slower response due to pH changes in the solution as hydrogen ions are released from the anode. These effects were also seen by Shiga and co-workers [136]. In a flowing system, the hydrogen ions are swept away and so only the fast response to ion migration in the gel remains. Advantages of this type of application are that large forces are not needed and the response time could be several minutes. 1.6.2

Light Modulators

A series of papers have described light modulators based on the temperature driven shrinkage of colored N-isopropylacrylamide in a second gel matrix [178, 183, 184]. As the colored particles expand at lower temperature, they extract water from the surrounding matrix and occupy most of the volume, thus cutting off light transmission (Figure 1.9). Clearly, an electrically driven system based on temperature or pH change could be constructed. Thermally responsive systems have also been developed for use as optical modulators. Particles of pigmented NIPAM gels are embedded in a dilute host gel that is selected to

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Figure 1.9 A gel-in-gel thermally driven light modulator at 20 °C (a, c) and 60 °C (b, d) (Reprinted by permission from Tsutsui, H., Mikami, M. and Akashi, R. All-polymer-gel light modulator consisting of a ‘gel-in-gel’ system, Adv. Mater., 16, 1925–9, Copyright (2004) Wiley-VCH Verlag GmbH).

allow ready exchange of water with the particles without opposing osmotic forces. The external gel must also be formed by cross-linking a polymer precursor in order to avoid forming an interpenetrating network with the gel particles. An external temperature increase results in a decrease in the absorbance of a thin sheet of the gel composite within one second as the particles shrink and the pigment becomes more localized [178, 183, 184]. 1.6.3

Gel Drug Delivery

Controlled drug release has been a subject of intense academic interest for many years. The bulk of commercial controlled release systems depend on pH sensitive release in the stomach or intestine. The matrix polymer may also dissolve or swell slowly to allow sustained release. Similar controlled-release systems may also be used to release flavors in food and rinks or to release fragrances during laundry or cleaning processes. Most such systems are essentially solid polymers because hydrogels would normally release the active small molecule too rapidly for most applications. This has changed in recent years as methods are developed to deliver much larger and sensitive protein and peptide drugs. These drugs also tend to be active at very low concentrations, so there is less need to produce high loadings of the drug in the matrix. Among the niche applications now being considered for hydrogels are the protection of protein drugs in the digestive system, adhesion of drug delivery patches to the mucus membranes and the protection of nanoparticulate drugs from the immune system [185].

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The fact that the swelling and contraction of these systems can be driven electrically, as discussed above, means that there are potential applications for adhesive patches [186] or implanted gels with electrically-triggered release. The thermo-responsive gels have been suggested for drug eluting stents, and here too an electrical mechanism might be used to trigger the temperature change and release [187]. 1.6.4

Gel Sensors

Many biochemical sensors depend on enzymes immobilized at an electrode surface, such that an analyte reacts to produce a compound that is readily oxidized or reduced at the electrode. Hydrogels bound to the electrode surface can be used to immobilize the enzyme without deactivating it [188]. In this case the gel is acting as a passive matrix but there are also systems where the responsiveness of the gel drives the sensing process. Gel-coated silicon microcantilevers can be used to detect swelling in response to changes in pH or other species [69, 189, 190]. Changes in swelling can also be sensed by attaching the gel to a pressure sensor. Sensors can also be based on fluorescence changes in sensitive molecules in the gel [191], on volume changes of gel particles changing the diffraction angle from a colloidal crystal [192, 193], by color changes in a gel hologram [194, 195] and other responses [196]. Sensors have been made which depend directly on the electroactivity of the gel. Glucose sensors have been made based on the binding of glucose by phenylboronic acid attached to a gel, where the change in ionic conductivity of the gel is measured as the binding reaction changes the swelling of the gel [197, 198]. Although the biocompatibility of gels is not an electroactive property, it should be noted that they may be important for implantable sensors. Many sensors have been developed that work well on implantation but almost all surfaces become coated with proteins and lose sensitivity over days after implantation [199]. Many gels are known to resist protein binding and may be able to protect implanted sensors and electrodes. A colloidal crystal contains an ordered array of sub-micron particles that will diffract light at an angle that depends on the spacing of the particles, following Bragg’s law. If the array is in a gel matrix, any volume change by the matrix will change the particle spacing and the diffraction angle, and so can be used as the basis of an optical sensor. This effect has been studied by many groups since the original work by Holtz and Asher [193, 192, 189].

1.7 Conclusions In just over 40 years since Katchalsky’s demonstration of a collagen engine, useful gel actuators remain elusive. A number of recent developments offer hope that this problem will be resolved before too long. Good chemically driven actuators can be made although they are not very practical for machinery. There have been significant advances in the understanding of gel responses to electrical fields and in the design of gels with superior mechanical properties. The challenge now is to design an electrochemical system that provides a suitable reversible chemical change to drive a gel and a mechanical design that

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subdivides the gel enough to give a rapid response. Problems such as this have been solved in the battery field many times, but with much greater inputs of researchers’ effort than have been applied to gels.

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2 Bio-Responsive Hydrogels for Biomedical Applications Tom McDonald1,2, Alison Patrick1,2, Richard Williams1,2, Brian G. Cousins1,2 and Rein V. Ulijn3 1

School of Materials, Materials Science Centre, University of Manchester, United Kingdom Manchester Interdisciplinary Biocentre (MIB), University of Manchester, United Kingdom 3 University of Strathclyde, United Kingdom

2

2.1 Introduction The term hydrogel is derived from the Latin prefix hydro- referring to water and gel, geluor gelatus- to describe a frozen or immobilised structure. Hydrogels are super absorbent materials, which consist of over 90–99 % water and are prepared by chemical polymerisation or physical assembly of man made or natural resources. Hydrogels form an important class of materials that impact upon modern medicine and are rapidly becoming significant due to the greater life expectancy that people will reach in the twenty-first century. Such materials have found applications such as intraocular and soft contact lenses, pharmaceutical carriers, drug delivery devices, biological sensors, wound dressings and scaffolds for regenerative medicine [1, 2]. Such materials are often nonfouling due to their high hydrophilicity, allow for facile diffusion of solutes and are easily functionalised to promote specific interactions with and respond to the biological environment [3]. In this chapter, the focus is on hydrogel materials that are designed to facilitate communication between the material and the biological environment. There are two main classes of hydrogels (Figure 2.1), those composed of threedimensional networks of cross-linked polymer chain structures that are insoluble in water (chemical hydrogels) and those produced by the self-assembly of (macro) molecules to form noncovalent structures (physical hydrogels).

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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(a) Chemical hydrogel

(b) Physical hydrogel

Figure 2.1 Schematic diagram of the two main classes of hydrogels produced by chemically cross-linked polymers (a) and (macro) molecules forming noncovalent structures (b).

2.2 Chemical Hydrogels Chemical hydrogels are composed of three-dimensional networks produced by the reaction of one or more (macro) monomeric structures through chemical cross-linking, resulting in insoluble structures that may be further stabilised by noncovalent interactions such as ionic, hydrophobic, pi-stacking, hydrogen bonding and/or van der Waals interaction [4]. The use of synthetic hydrogels dates back to the 1950s when Wichterle and Lim synthesised the first hydrogels for biomedical applications to develop soft contact lenses based on a copolymer of 2-hydroxyethyl methacrylate and ethylene dimethacrylate to form poly(hydroxyethyl methacrylate) (PHEMA) [5]. The commercial success of contact lens materials stimulated vast interest in hydrogels and, more recently, led to the development of stimuli-responsive or ‘smart’ hydrogels that change their physical properties in response to changes in the local environment such as pH, temperature, ionic strength, solvent composition, pressure and electrical potential. Responsive chemical hydrogels typically change their hydration states, that is they swell or collapse in response to applied stimuli.

2.3 Physical Hydrogels Physical hydrogels are composed of molecular building blocks varying from small amphiphilic molecules to macromolecules, which form by self-assembly, exclusively through noncovalent cross-links. Water can penetrate throughout the physical structure in between molecular building blocks, which gives rise to hydrogel network structures. Physical hydrogels typically show gelation or dissolution behaviour in response to applied stimuli.

2.4 Defining Bio-Responsive Hydrogels By using the similar design principles for both hydrogel types, they can be tailored to incorporate recognition motifs that respond to specific biomolecular events.

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Here, the term ‘bio-responsive’ is defined as stimuli responsive materials that change properties in response to a biomolecular recognition event, whereby these molecular interactions are translated into bulk changes in material properties that result in an observed response, which may be optical, electronic or chemical in nature (Figure 2.2).

Input

Design

Step 1: Stimulus The biochemical stimulus may be a small molecule (peptide, glucose, inorganic ions) or a biomacromolecule (antibody, enzyme).

Step 2: Molecular Biorecognition Recognition of the stimulus by the incorporated receptor/sensing element.

Step 3: Molecular Actuation A molecular change triggers a macroscopic transition such as collapse, swelling, selfassembly.

Step 4: Response The macroscopic transition causes emission of an optical, electronic or chemical signal. Output

Figure 2.2 A simplified scheme of events that define and influence a bio-responsive hydrogel system. An input of a stimulus (step 1) leads to the binding of the receptor to the ligand within the gel network resulting in molecular biorecognition events (step 2). Molecular actuation is stimulated in step 3 to trigger further micro and macroscopic events, i.e. swelling. Molecular actuation results in a response (step 4) that can be mechanical, chemical, optical or electronic.

Indeed, bio-responsive hydrogels have gained significant interest over the last several years for applications related to drug delivery, diagnostics, wound healing and tissue regeneration. This exciting field is vast and the emergence of new articles appears on a daily basis and continues to increase rapidly. The authors have therefore chosen a relatively small number of studies from the literature to outline each example in detail. The reader is encouraged to locate further examples of bio-responsive systems cited elsewhere in the literature [1, 3, 6, 7].

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Biomedical Applications of Electroactive Polymer Actuators

2.5 Bio-Responsive Chemical Hydrogels This section looks at three examples of actuation for bio-responsive chemical hydrogels and describes how these methods have been used in the development of biosensor, drug delivery and tissue scaffold systems. These three actuation examples are classified in Table 2.1 as follows: changes in cross-linking density (1–8), electrostatic interactions (9–13) and molecular conformation (14, 15). 2.5.1

Actuation Based on Changing the Cross-Linking Density

Cross-linking density may be modified by either the cleavage of polymer chains (irreversible) or via competition at the binding sites of the cross-linking moieties (reversible). An antigen-responsive hydrogel has been described by Uragami and co-workers [8], (Table 2.1, entry 1). This semi-interpenetrating network (semi-IPN) hydrogel contains both grafted antigens and their corresponding specific antibody. Antigen–antibody binding causes the formation of cross-links within the polymer. Free antigens in the solution around the hydrogel create competition between the grafted antigens leading to a decrease in crosslinking density, thus an increase in swelling of the hydrogel. This increase in swelling is reversible. By removing the hydrogel from the antigen solution and washing, the polymer would return to approximately its original volume. This response also corresponds to an increase in permeability of the polymer, and by using the polymer as a membrane, selective permeation of a protein was shown in response to the specific antibody (Figure 2.3). More recently a similar design was employed to produce an antigen-responsive membrane that has gating properties for selective diffusion in response to the presence of a free antigen [9] (Table 2.1, entry 2). A glucose responsive hydrogel has been developed by using the formation of crosslinking complexes between poly(glucosyloxy-ethy1methacrylate) poly(GEMA) and concanavalin A (Con A) (Table 2.1, entry 3) [10]. When there is free glucose present the hydrogel responds through an increase in swelling. Systems based on the cleavage of polymer chains as the mechanism for changing crosslinking density include Table 2.1, entries 4–6: a chymotrypsin-responsive hydrogel developed by Moore and co-workers [11], a cell-responsive hydrogel demonstrated by Hubbell’s group in which enzymes secreted by the cells cleave cross-links within the polymer [12], and also a calcium-responsive hydrogel by Golbart and Kost. Here an inactive form of the enzyme is activated by calcium ions from solution leading to the degradation of the polymer matrix [13]. Messersmith and co-workers have developed a system in which the assembly of short peptides is controlled by the formation of chemical cross-links. The stimulus is the crosslinking enzyme, Transglutamase. Eleven peptide residues of alternating hydrophobic/ hydrophillic charge are used. The biorecognition sites are intermolecular glutamine and lysine residues. The molecular activation is the formation of covalent cross-links, which stabilise the peptides into fibrils that provide the response, a hydrogel [14] (Table 2.1, entry 7). In a second system from Lutolf’s group, the lysine donor is connected by a peptide sequence, GPQGIWGQ, a biorecognition motif for the cellularly excreted stimulus, the enzyme MMP-1 [15] (Table 2.1, entry 7). The molecular response in this case is the enzymatic hydrolysis of peptide cross-links. By incorporation of peptides that can be

Bio-Responsive Hydrogels for Biomedical Applications

47

Table 2.1 Examples of some bio-responsive hydrogels described in the literature

1 2 3 4 5 6 7 8

9

Stimulus

Actuation

Antigen

Cross-linking: Antigen– antibody binding

Hydrogel

Response

Poly-acrylamide Increased swelling (reversible) Antigen Cross-linking: Antigen– Dextran Gating of antibody binding polymeric membrane network (reversible) Glucose Cross-linking: Poly(GEMA) Increased Con A– glucose binding swelling (reversible) Enzyme Cross-linking: Cleavage Poly-acrylamide Gel dissolution (Chymotrypsin) (irreversible) Enzyme (MMPs) Cross-linking: Cleavage PEG Gel dissolution (irreversible) Cellulose-Starch Degradation of Calcium Cross-linking: Matrix matrix degradation (low cross-link (irreversible) density) Cross-linking: Cleavage of Peptide Hydrogel Enzyme hydrogel formation (Transglutaminase/ strands MMP-1) PNIPAM-coChange in Biotin Cross-linking: focus of image Displacement of molecules AAC projected through material Glucose

10 Glucose

11 Enzyme (Protease)

12 Enzyme (Protease)

13 Specific ligand 14 Specific ligand

Electrostatic: GOx oxidation of glucose causing pH responsive swelling Electrostatic: GOx oxidation of glucose causing pH responsive collapse of gating polymer Electrostatic: Protease cleavage of zwitterions resulting in a net charge and swelling Electrostatic: Same as entry 11 except net charge upon zwitterion cleavage is the same as the charge of the protein to be released Change in conformation: Incorporated protein Change in conformation: Incorporated protein

Ref. [8] [9] [10] [11] [12] [13] [14, 15] [16]

Poly (HEMA-co- Increased swelling DMAEMA (reversible)

[18]

Membrane of PVDF with grafted PAAC

Gating of membrane (reversible)

[19]

PEGA

Increased swelling (irreversible)

[20]

PEGA

Increased swelling (irreversible)

[22]

PEG and the protein CaM

Increased swelling (reversible) Increased swelling (reversible)

[24]

PEG and the protein CaM

[25]

48

Biomedical Applications of Electroactive Polymer Actuators (a)

H O AAm

N C CH

APS/ TEMED

Modified antibody H O N C CH

Antigen

Polymerized antibody (c)

CH2

Modified antigen (δ ), AAm MBAA, APS/TEMED

Equilibrium swelling ratio (m3/m3)

(b)

1.15 Antigen-antibody semi-IPN hydrogel

Antigen

1.10

1.05

1.00

0.95

PAAm semi-IPN hydrogel

0

2

4

6

8

10

Antigen concentration (mg ml–1)

Figure 2.3 Antigen-responsive hydrogel. (a): Synthesis of the antigen–antibody semi-IPN hydrogel. (b): Effects of the free antigen concentration on the hydrogel swelling ratio. (c): Diagram of a suggested mechanism for the swelling of an antigen–antibody semi-IPN hydrogel in response to a free antigen. (Reprinted with permission from [8]. Copyright (1999) Nature Publishing Group.)

cleaved by cell-secreted proteases, cells can be encouraged to migrate through a gel by digesting it locally, similar to cell migration in natural extracellular matrices. By including these motifs in artificial matrix mimics, it improves the biological compatibility. Cells can be cultured on artificial matrix, and then begin to replace the artificial matrix with a natural one. In another example of actuation by a change in cross-linking density, the optical properties of spherical hydrogel particles are exploited as flexible microscopic lenses that change their focus upon changes in swelling (Table 2.1, entry 8). Kim and co-workers [16] developed a system in which a reaction between the biomolecule and cross-links in the hydrogel structure leads to swelling at the surface of the material. Hydrogels of poly(N-isopropylacrylamide-co-acrylic acid), P(NIPAM-co-AAc), were prepared and functionalised with biotin and aminobensophenone. These hydrogels were adsorbed onto a surface. Anti-biotin molecules were photochemically cross-linked into the structure at the surface, linking the biotin and aminobenzophenone. This cross-linking controls the degree of swelling at the surface. When free biocytin (substitute for biotin) is present in solution it displaces the anti-biotin from the anchored hydrogel biotin, breaking the cross-links and changing the degree of swelling at the surface, as shown in Figure 2.4. The degree of swelling at the hydrogel surface controls the optical properties. Actuation results in a

Bio-Responsive Hydrogels for Biomedical Applications

49

change in the focus of an image projected through the material, observed by optical microscopy. This can be used as the output method to determine that a change in swelling has occurred. Some results are shown in Figures 2.4c and 2.4d before and after treatment with biocytin respectively. Before treatment the hydrogel focuses the image so a white edge and dark centre is seen. After treatment the hydrogel de-focuses and a white centre is seen.

P(NIPAM-co-AAc) H2 C

H C C

H2 C n O

NH

uv

H C C

m O

(b)

OH

CH H3C

(c)

(a)

CH3

Biotin

Aminobenzophenone

Anti-Biotin

(d)

Biocytin

Cross-linked area

Figure 2.4 (a) Structure of P(NIPAM-co-AAc). (b) Schematic diagram showing hydrogels functionalised with biotin and aminobenzophenone, adsorbed to a surface. Anti-biotin crosslinks are formed to control degree of swelling at the surface. Optical microscopy images: (c) before treatment with biocytin, (d) after treatment with biocytin (scale bar ¼ 2 mm).

Other bio-responsive hydrogels have been developed to give an optical output that use different actuation methods. These include sensors for a-cyclodextrin, in which a reaction with polydiacetylene liposomes embedded in poly(ethylene glycol) diacrylate (PEGDA) leads to a visible colour change of the hydrogel from blue to red [17]. 2.5.2

Actuation Based on Changes in Electrostatic Interactions

A second mechanism to actuate a change in swelling is through electrostatic repulsion/ attraction of polymer chains. This method of actuation is widely studied in the context of glucose responsive polymers for the treatment of diabetes. Generally, glucose oxidase (GOx) is immobilised in the hydrogel, along with catalase, an enzyme that is used to prevent the build up of hydrogen peroxide. When glucose is present it is converted to gluconic acid and hydrogen peroxide, which lowers the pH within the hydrogel. There are two different designs in which this decrease in pH is used to actuate a change in swelling: matrix type systems, where the enzyme and insulin are contained within a bulk polymer, or membrane type systems, where the drug is contained in a reservoir within a membrane. Kost and co-workers produced an example of the matrix system where the insulin and enzyme were contained uniformly throughout a solid polymer [18] (Table 2.1, entry 9). This polymer contained amine groups and had a low cross-linking density, so that at low pH the amines become ionised leading to an increase in swelling. In Figure 2.5, GOx catalyses the reaction of glucose to gluconic acid forming hydrogen peroxide (equation 1). A build up of hydrogen peroxide leads to inhibition of the enzyme, and because oxygen is needed to form gluconic acid a shortage of oxygen leads to slower

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Biomedical Applications of Electroactive Polymer Actuators

swelling rates. For this reason this system also contained catalase, which serves to convert hydrogen peroxide to water and oxygen (equation 2). The device was implanted in rats and these in vivo experiments indicated that some of the entrapped insulin retained its active form and was effective in reducing blood glucose levels. A membrane type system has been created by Liang and co-workers [19] (Table 2.1, entry 10). Here, poly(acrylic acid) (PAAC) was grafted to a porous membrane of poly(vinylidene fluoride) (PVDF) and GOx was covalently bound to the PAAC. The PAAC chains ‘gate’ the membrane pores and in the presence of glucose the chains collapse and the pores are opened.

GOx Glucose + O2 + H2O

2H2O2

Figure 2.5

→

Gluconic acid + H2O2

Catalase → O2 + 2H2O

(1)

(2)

Reactions driven by the enzymes used in glucose responsive hydrogels.

A different type of electrostatically actuated responsive hydrogel was developed by Ulijn and co-workers [20] (Table 2.1, entry 11). The stimulus in this system is an enzyme, specifically a protease (an enzyme that hydrolyses peptide bonds). Enzymes are well suited to use as a stimulus as they can be targeted to highly specific substrates thanks to unique chemo-, regio- and enantio-selective mechanisms [21]. They work under aqueous physiological conditions, that is high ionic strength, pH 5–8 and at 37 °C. These systems are easily matched to applications through a wide range of functions and perform key roles as selective catalysts in cell pathways and disease states. Here the response is governed by a peptide actuator which consists of two components: a neutral dipeptide flanked by oppositely charged actuating amino acids, creating an overall neutral, zwitterionic peptide chain. The central peptide is the enzyme cleavable linker (ECL), which can be matched to be cleaved exclusively by a target protease. When hydrolysis occurs, a mobile anionic peptide fragment is cleaved, leaving a cationic peptide fragment attached to the hydrogel. Electrostatic repulsion between these groups leads to an increase in swelling, causing the entrapped payload to diffuse out of the hydrogel into the surrounding solution. More recently, this system has been further developed to release proteins [22] (Table 2.1, entry 12). This was achieved by tailoring the design of the peptide actuator to give a net charge after ECL cleavage that is matched to the charge on the protein. Electrostatic repulsion then assists the release of the protein. A biosensor system that uses such ECL recognition followed by actuation has been developed for the detection of elastase. It uses fluorescence resonance energy transfer (FRET) as a method of displaying the actuation [23] (Table 2.1, entry 14). In this biosensor example a FRET pair is incorporated into the hydrogel, poly(ethylene glycol acrylamide) (PEGA), and is separated by an ECL. Enzyme action cleaves the ECL, removing the acceptor molecule of the FRET pair, switching on fluorescence. These hydrogels are also functionalised with negatively charged residues that cause electrostatic repulsion and lead to increased swelling of the hydrogel for easier diffusion of the enzyme into the particle. As elastase is positively charged at neutral pH, the negative charges added into the hydrogel are bifunctional and attract the elastase to the ECL once inside the hydrogel (Figure 2.6).

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Figure 2.6 (a) Peptide designed for the release of negatively charged protein molecules. Two positively charged amino acids (arginine) are separated from two negative amino acids (aspartic acid) by the neutral ECL (two alanine residues). A single net negative charge remains on the particle following enzyme hydrolysis. (b) Exclusion of albumin from the negatively charged swollen particle occurs following hydrolysis of the bond between alanine residues by thermolysin. (c) Two photon microscopy images of Texas Red labelled albumin being released from the particles (scale bars ¼ 75 mm). (Reprinted with permission from [22]. Copyright (2008) Royal Society of Chemistry.)

2.5.3

Actuation Based on Conformational Changes

The third mechanism of actuation is based on changes in conformation of natural proteins (Table 2.1, entries 13 and 14). These systems incorporate a natural protein into the hydrogel that undergoes a conformation change and thus alters the characteristics of the material. The use of proteins as actuators is a new development in bio-responsive hydrogels. An example of this system was developed by Mrksich and co-workers [24]. In this study the functional nature of the hydrogel was given by the conformational properties of the protein calmodulin (CaM). Calmodulin has two distinct conformational states. In the presence of calcium ions, CaM has an extended, dumb-bell shaped conformation (extended CaM).

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Biomedical Applications of Electroactive Polymer Actuators

Figure 2.7 Ligand-responsive hydrogel that relies on conformational changes. (a) The two conformational states of CaM: an extended conformation in the presence of calcium ions (left) and a collapsed conformation upon binding to a ligand (right). (b) Shown is a hydrogel with CaM in a ligand-free state (left) and the same gel with CaM in a ligand-bound state (right) (scale bars ¼ 1 mm). (c) Hydrogels were exposed to TFP ligand and the volume was measured at various intervals for two hours. The gel was then washed repeatedly and incubated in a calcium-containing buffer to restore the extended CaM conformation. (Reprinted with permission from [27]. Copyright (2006) American Chemical Society.)

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This calcium-bound extended CaM undergoes a transition from an extended dumb-bell to a collapsed conformation (collapsed CaM) upon the binding of ligands. The CaM protein is incorporated into the hydrogel by covalently reacting it with four armed PEG molecules. The hydrogel shows a macroscopic decrease in volume when exposed to the trifluoperazine ligand (TFP). TFP binds specifically to CaM, causing the CaM to undergo a conformation change from extended to collapsed (Figure 2.7). This decrease in volume could be reversed by chelating the calcium ions, thus removing the calcium-bound ligand. Numerous cycles between extended and collapsed material were possible, demonstrating the reversible nature of the hydrogel. More recently, this concept has been developed to incorporate a photochemical assembly, allowing spatial control of the location of dynamic proteins [25].

2.6 Bio-Responsive Physical Hydrogels The spontaneous assembly of (macro) molecules to form physical hydrogels is an area of increasing importance. It is desirable to introduce a level of control in these systems, enabling the triggering of hydrogel formation or dissolution upon application of selective stimuli that are uniquely found in certain biological situations. This can be achieved by employing biorecognition elements into (some of) the molecular building blocks in the direct vicinity of ionisable groups, or hydrophobic regions. Stimuli that can be exploited in this context include altering of physical conditions, that is temperature, pH, ionic strength and oxidative species, or using enzymes, such as proteases, kinases, phosphatases or ligases. Enzymatic actuation leads to controlled intramolecular events, such as shortening/straightening of a peptide chain, ionisation, side chain modification or the exposure of hydrophobic regions, to redefine existing molecular interactions that enable new intermolecular associations [21]. The next section gives examples of mechanisms that may be used in physical gels based on small molecules or macromolecules. The biorecognition elements are variable depending upon enzymatic function, for example a peptide bond between specific amino acids (Table 2.2, entries 3, 4, 7–9), a (phosphate) ester group (entries 1–3, 5) or a specific sequence of DNA (entry 6). 2.6.1

Enzyme-Responsive Physical Hydrogels

The formation of physical assemblies from small peptide amphiphiles can be achieved by manipulating the structure of the peptides with enzymes. These reactions are typically the formation/breaking of covalent bonds altering the structures and thereby controlling the magnitude of weak structural forces, typically p–p interactions, hydrogen bonding and electrostatic interactions. By appropriate choice of pairs of enzymes, hydrogelation can be made reversible [26]. Typically, self-assembly is a one-way process, as the enzyme only operates in one direction governed by thermodynamic equilibrium, but by using two enzymes with opposite actions, one under thermodynamic control and the other involving a coupled reaction, this limitation may be overcome. For example, peptide (de-) phosphorylation can be exploited as an effective means of hydrogelation control, where a kinase in the presence of ATP (a natural source of energy) adds the phosphate group, while a phosphatase cleaves it off (Figure 2.8a and 2.8b). The phosphate group is used to prevent self-assembly of beta sheet forming

54

Biomedical Applications of Electroactive Polymer Actuators

Table 2.2 Entries 1–7 highlight mechanism of physical hydrogel formation. Entries 8 and 9 show self-assembly directed through the formation/disruption of covalent cross-links Stimulus

Biorecognition Molecular Actuation

Response

Ref.

RGYSLG

Soluble b-sheets* Hydrogel formation* Hydrogel formation* Hydrogel formation Hydrogel formation Hydrogel formation Drug release Drug release Cell migration

[29]

1 Phosphatase/ Kinase 2 Phosphatase/ Kinase 3 Thermolysin/ Subtilisin 4 Thermolysin

Fmoc-X + FF

5 Penicillin G Amidase 6 T4 DNA ligase

Phenylacetic acid Base pairs

Triple elimination

7 Chymotrypsin 8 Urokinase 9 MMP-2

FA SGRSANA GTAGLIGQ

Bond cleavage b-sheet disruption Hydrophobic disruption

Y/ phosphate Y + L/F-OMe

Phosphate group cleavage/addition Phosphate group cleavage/addition Peptide synthesis/ ester cleavage Peptide synthesis

DNA ligation

[27, 30–32] [33] [26] [34] [28] [35] [36] [37]

Letters indicate single letter code for amino acids as follows: A ¼ Alanine, F ¼ Phenylalanine, G ¼ Glycine, I ¼ Isoleucine, K ¼ Lysine, N ¼ Asparagine, P ¼ Proline, Q ¼ Glutamine, R ¼ Arginine, S ¼ Serine, T ¼ Threonine W ¼ Tryptophan, X ¼ any amino acid and Y ¼ Tyrosine. * Indicates reversible two-enzyme system.

peptides or small aromatic peptide derivatives, typically 1–5 amino acids long. This system is used for a range of applications: sensing the presence of enzymatic inhibitors, the signal being prevention of gelation [30]; the reversible formation of an extracellular matrix mimic for cell culture in vivo [27]; and an antimicrobial system, in which the self assembly is triggered inside microbes using their own enzymes, killing the cell [31]. Table 2.2, entry 4 is a system based on subtilisin and thermolysin [33]. The protease thermolysin acts as the stimulus for assembly, operating to form dipeptide derivatives from Fmoc-threonine and either a leucine or a phenylalanine methyl ester (Figure 2.8c). The dipeptide derivatives formed give rise to self assembled nanofibrous hydrogels. The hydrogel can be dissolved by employing a second enzyme, subtilisin, which cleaves the methyl ester to leave a soluble Fmoc-dipeptide. This system could be used to entrap cells in a network using one enzyme, and release them upon application of the other. A system for controlled release of a payload from a physical gel is discussed in Table 2.2, entry 8 [36]. By encasing a drug within a network which can be degraded by an enzyme, a responsive system has been created – b-sheet segments (KLD)12 either side of an enzyme cleavable biorecognition motif, SGRSANA. A drug is linked to a similar sequence, and when mixed in aqueous solvent, form a gel through stacking of the b-sheets. The stimulus is the addition of the protease urokinase plasminogen activator. This acts from the outside of the gel, through cleavage of the biorecognition motif. Once cleaved, the resulting molecular activation is a weakening of the interactions and breaking down the self-assembled structure. The response to the stimulus is the gradual release of the encapsulated drug. Enzyme-responsive hydrogels have also been studied as extracellular matrix mimics. For example, Hartgerink and co-workers investigated the use of peptide ampiphiles which self-assemble into fibrous cylindrical micelles driven by the interactions of hydrophobic

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Figure 2.8 Reversible self-assembly by the use of enzymes. (a) De-phosphorylation of the target molecule results in the formation of a self-assembled network. By introducing an inhibitor to the system, hydrogel formation is inhibited (Reprinted with permission from Yang, Z and Xu, B. Bio-responsive hydrogels for biomedical applications, Chem. Commun., 21, 2424–25, Copyright (2004) Royal Society of Chemistry). (b) A reversible system of assembly by dephosphorylating a peptide with phosphatase to trigger assembly, and reversing the assembly with the use of a kinase (in the presence of ATP) (Reprinted in part with permission from Yang, Z., Liang, G., Wang, L. and Xu, B. Using a Kinase/Phosphate Switch to Regulate a Supramolecular Hydrogel and Forming the Supramolecular Hydrogel in vivo, J. Am. Chem. Soc., 128, 9, 3038–43, Copyright (2006) American Chemical Society). (c) A reversible system in which the formation of a peptide bond results in an assembling Fmoc-dipeptide methyl ester. Assembly is reversed by the use of subtilisin to promote ester hydrolysis (Reproduced with permission from Das, K. A., Collins, R. J. and Ulijn R. V., Exploiting Enzymatic (Reversed) Hydrolysis in Directed Self-Assembly of Peptide Nanostructures, Small, 4, 279–87. Copyright (2008) Wiley-VCH Verlag GmbH).

tails and hydrophilic heads with water (Table 2.2, entry 9) [37]. These peptides contain RGD, a cell binding epitope, which is displayed on the surface allowing their use as a cell culture medium. To allow the naturally secreted enzyme MMP-2 to act as a stimulus for matrix remodelling, the biorecognition motif GTAGLIGQ is inserted between the tail and the head. The molecular response to the stimulus is the disruption of the hydrophobic effects and breaking down the fibres [38]. The response to the stimulus was that the cells could be observed migrating along the fibres.

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Biomedical Applications of Electroactive Polymer Actuators

2.7 Electroactive Chemical Hydrogels Electroactive chemically cross-linked hydrogels integrate both polymer hydrogels and conjugated (conducting) electroactive polymers, that is forming an electroactive hydrogel. In these systems, a biochemical reaction results in a current which can easily be read. This type of system is especially useful for observing the action of redox enzymes. An example of this combination of polymers is the work done by Brahim and co-workers [39] on a glucose sensing system, in which platinum electrodes were coated with cross-linked poly(hydroxyethyl methacrylate) (pHEMA) and polypyrrole components and contained GOx entrapped in the matrix (Figure 2.9). When the glucose is present it reacts with the enzyme and produces hydrogen peroxide. The steady state current produced by this reaction can be measured and gives a reading of the enzyme activity which is dependent on the enzyme substrate concentration. The pHEMA component creates a biocompatible environment for the entrapped enzyme, giving the sensor a lifetime of up to one year. Brahim and coworkers have also shown in this work that by switching the polypyrrole component for dimethylaminoethyl methacrylate (DMA) the system can release a drug upon stimulation, becoming a drug delivery system rather than a sensing system [40].

Figure 2.9 Schematic diagram showing platinum electrodes coated with cross-linked pHEMA and polypyrrole components with GOx entrapped to form glucose sensor that displays result as current.

This is an example of a biosensor that uses current as an output. Other biosensor systems have been developed that use similar methods for the detection of cholesterol, glucose and hydrogen peroxidise [41, 42].

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2.8 Conclusion In summary, it has been demonstrated that bio-responsive hydrogels allow effective and selective response to a variety of biochemical stimuli, usually by displaying a swelling/ collapse or gelation/dissolution response. While many systems have been studied in well defined conditions, future development should allow devices to be created that can actively and rapidly respond to biological stimuli at physiological levels and in complex biological fluids. Due to lengthy regulatory and approval processes, it is likely that applications will focus on the use of US Food and Drug Administration (FDA) approved polymers, such as pHEMA, PEG and PGLA based hydrogels. In physical gels, the use of a wide range of enzymes to control gelation has been demonstrated. Many of these systems are at the proof of concept stage, but increasing numbers of examples of biological relevance have been demonstrated, typically in the realm of cell culture. Increasingly, the ability to selectively control the formation and growth of hydrogels will be shown to be of tremendous value for a range of applications, such as in vivo tissue repair, biosensing and drug delivery.

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15. Ehrbar, M., Rizzi, S. C., Schoenmakers, R. G., et al. (2007) Biomolecular hydrogels formed and degraded via site-specific enzymatic reactions, Biomacromolecules, 8, 3000–7. 16. Kim, J. S., Singh, N. and Lyon, L. A. (2007) Displacement-induced switching rates of bioresponsive hydrogel microlenses, Chem. Mat., 19, 2527–32. 17. Lee, N. Y., Jung, Y. K. and Park, H. G. (2006) On-chip colorimetric biosensor based on polydiacetylene (PDA) embedded in photopolymerized poly(ethylene glycol) diacrylate (PEG-DA) hydrogel, Biochem. Eng. J., 29, 103. 18. Traitel, T., Cohen, Y. and Kost, J. (2000) Characterization of glucose-sensitive insulin release systems in simulated in vivo conditions, Biomat., 21, 1679–87. 19. Chu, L. Y., Li, Y., Zhu, J. H., et al. (2004) Control of pore size and permeability of a glucoseresponsive gating membrane for insulin delivery, J. Controlled Release, 97, 43–53. 20. Thornton, P. D. M., Mart, R. J. and Ulijn, R. V. (2007) Enzyme-responsive polymer hydrogel particles for controlled release, Adv. Mat., 19, 1252–6. 21. Ulijn, R. V. (2006) Enzyme-responsive materials: a new class of smart biomaterials, J. Mat. Chem., 16, 2217–25. 22. Thornton, P. D. M., Mart, R. J., Webb, S. J. and Ulijn, R. V. (2008) Enzyme-responsive hydrogel particles for the controlled release of proteins: designing peptide actuators to match payload, Soft Matter, 4, 821–7. 23. Patrick, A. G. and Ulijn, R. V. (2007) Fluorescent hydrogel sensor particles for detection of elastase, Mat. Res. Soc. Proc., 1063–PP06–05. 24. Murphy, W. L. D., Dillmore, W. S., Modica, J. and Mrksich, M. (2007) Dynamic hydrogels: translating a protein conformational change into macroscopic motion, Angew. Chem. Int. Edn., 46, 3066–9. 25. Sui, Z., King, W. J. and Murphy, W. L. (2007) Dynamic materials based on a protein conformational change, Adv. Mat., 19, 3377–80. 26. Toledano, S., Williams, R. J., Jayawarna, V., and Ulijn, R. V. (2006) Enzyme-triggered selfassembly of peptide hydrogels via reversed hydrolysis, J. Am. Chem. Soc., 128, 1070–1. 27. Yang, Z. M., Liang, G. L., Wang, L. and Bing, X. (2006) Using a kinase/phosphatase switch to regulate a supramolecular hydrogel and forming the supramoleclar hydrogel in vivo, J. Am. Chem. Soc., 128, 3038–43. 28. Um, S. H., Lee, J. B., Park, N., et al. (2006) Enzyme-catalysed assembly of DNA hydrogel, Nature Mat., 5, 797–801. 29. Winkler, S., Wilson, D. and Kaplan, D. L. (2000) Controlling beta-sheet assembly in genetically engineered silk by enzymatic phosphorylation/dephosphorylation, Biochem., 39, 12739–46. 30. Yang, Z. M. and Xu, B. (2004) A simple visual assay based on small molecule hydrogels for detecting inhibitors of enzymes, Chem. Comm., 2424–5. 31. Yang, Z., Liang, G., Guo, Z. and Xu, B. (2007) Intracellular hydrogelation of small molecules inhibits bacterial growth, Angew. Chem. Int. Edn., 46, 8216–9. 32. Yang, Z. M., Liang, G. L., Ma, M. L., et al. (2007) In vitro and in vivo enzymatic formation of supramolecular hydrogels based on self-assembled nanofibers of a beta-amino acid derivative, Small, 3, 558–62. 33. Das, A. K., Collins, R., Ulijn, R. V. (2008) Exploiting enzymatic (reversed) hydrolysis in directed self-assembly of peptide nanostructures, Small, 4, 279–87. 34. Alder-Abramovich, L., Perry, R., Sagi, A., et al. (2007) Controlled assembly of peptide nanotubes triggered by enzymatic activation of self-immolative dendrimers, Chembiochem., 8, 859–62. 35. van Bommel, K. J. C., Stuart, M. C. A., Feringa, B. L. and van Esch, J. (2005) Two-stage enzyme mediated drug release from LMWG hydrogels, Organic and Biomol. Chem., 3, 2917–20. 36. Law, B., Weissleder, R. and Tung, C. H. (2006) Peptide-based biomaterials for proteaseenhanced drug delivery, Biomacromolecules., 7, 1261–5.

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37. Jun, H. W., Yuwono, V., Paramonov, S. E. and Hartgerink, J. D. (2005) Enzyme-mediated degradation of peptide-amphiphile nanofiber networks, Adv. Mat., 17, 2612–7. 38. Yang, Z., Liang, G. and Xu, B. (2008) Enzymatic hydrogelation of small molecules, Acc. Chem. Res., 41, 315–26. 39. Brahim, S., Narinesingh, D. and Guiseppi-Elie, A. (2002) Bio-smart hydrogels: co-joined molecular recognition and signal transduction in biosensor fabrication and drug delivery, Biosensors & Bioelectronics, 17, 973–81. 40. Brahim, S., Narinesingh, D. and Guiseppi-Elie, A. (2002) Bio-smart materials: Kinetics of immobilised enzymes in p(HEMA)/p(pyrrole) hydrogels in amperometric biosensors, Macromolecular Symposia, 186, 63–73. 41. Ivekovic, D., Milardovic, S. and Grabaric, B. S. (2004) Palladium hexacyanoferrate hydrogel as a novel and simple enzyme immobilization matrix for amperometric biosensors, Biosensors and Bioelectronics., 20, 872–8. 42. Sun, Y. X., Zhang, H. T., Huang, S. W. and Wang, S. F. (2007) Hydrogen peroxide biosensor based on the bioelectrocatalysis of horseradish peroxidise incorporated in a new hydrogel film, Sensors and Actuators B-Chem., 124, 494–500.

3 Stimuli-Responsive and ‘Active’ Polymers in Drug Delivery Aram Omer Saeed1, Jo´hannes Pa´ll Magnu´sson1, Beverley Twaites2 and Cameron Alexander1 2

1 School of Pharmacy, University of Nottingham, UK School of Pharmacy and Biomedical Sciences, University of Portsmouth, UK

3.1 Introduction The efficacy of synthetic drug compounds in therapy is strongly dependent on their formulation into medicines and on their distribution, localisation and accumulation in different regions in the body. The ability to deliver a drug compound to the specific target site (organ, tissue, cell, intracellular compartment) remains a challenge. A century on from Paul Ehrlich’s visionary hypothesis, for which he received the Nobel Prize in 1908, there are still no ‘Magic Bullets’ [1] against most diseases. As a result, the delivery of drugs to the desired target sites in the body at the right time and in the right dose is still an unmet clinical need [2]. This leads to inefficient use of drugs, undesired side effects and greater medical intervention, with resulting burdens on patients, carer populations and healthcare budgets. Current targeting methods for many drugs either lack specificity or are not active for certain patient groups. As a consequence, drug delivery systems are gaining in importance – and many of these are based on polymers [2, 3]. In addition, new generations of therapeutics are emerging, and these too are often macromolecular or polymeric in nature. For example, in 2002 and 2003 the US Food and Drug Administration (FDA) approved more biotechnology products (proteins and antibodies) and drug delivery systems as marketed products than new low molecular weight drugs. However, these new classes of drugs and their conjugates, complexes and formulated vehicles – sometimes considered as Nanomedicines – and the related biotherapeutics, still urgently require technologies to

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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ensure their specific localisation in target sites [4, 5]. As a consequence, new modalities of ‘smart’ or active materials ranging from bioresponsive to electroactive polymer–drug delivery systems are being developed. In this chapter, a short introduction is given to key concepts in drug delivery, the specific issues for polymers and smart materials related to the transport and release of therapeutic compounds are considered, and some exciting recent examples where actuating polymers have been developed to help target and deliver drug molecules are highlighted. The focus is primarily on soluble polymers and nanoparticles rather than hydrogels, as there are already numerous excellent reviews on responsive hydrogels available, [6–16] and the design concepts in the soluble and nanoparticle polymer systems best exemplify the key responses that can be engineered into ‘smart’ materials.

3.2 Drug Delivery: Examples, Challenges and Opportunities for Polymers In general, drug delivery aims at optimising therapy by delivering bioactive agents at specific sites or specific rates to the patient. The field has evolved from simple topical waxes and delayed release formulations to the targeted delivery of therapeutic agents to specific cells and subcellular compartments. Traditional approaches to drug delivery have been based on simple formulation parameters. For low molar mass compounds (Mw <1000), which still represent the main focus for the pharmaceutical industry, the goal is to deliver the drug to the pharmacological receptor or biochemical target with as little complexity in the formulation as possible. Since many of these compounds are relatively lipophilic and can partition across membranes, the drug delivery system needs to address the biopharmaceutical requirements of the dosing route, and these involve a variety of physical and biochemical barriers. Of course, delivery is ultimately dependent both on the physicochemical properties of the drug and on the physiology of the body, thus a brief consideration of these factors and the common dosing routes is needed to put polymer-mediated drug delivery in context. 3.2.1

Oral Drug Delivery Systems

The oral route is the most widely used for drug administration and is likely to remain the most favoured route in the future owing to good patient compliance, familiarity and long history of successful development in the pharmaceutical industry. However, as many candidate therapeutics in development are increasingly of higher molar mass and poorer aqueous solubility compared to existing drugs, the systems for formulating and delivering these candidates are having to become more complex. For conventional oral delivery water-soluble drugs can be formulated into tablets or capsules with controllable erosion or breakdown profiles to determine rate and site of drug release in the gastrointestinal (GI) tract. The types of polymers used in these systems are relatively simple, and include cellulose and derivatives, for example 2-hydroxypropylmethylcellulose (HPMC) and sodium carboxymethylcellulose, which are often combined with other materials such as poly(N-vinylpyrollidone) and poly(methacrylic acid/methacrylates) for optimal formulation properties [17]. There are a great number of potential variations possible in these drug delivery materials and many have been successfully commercialised.

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However, in the case of many poorly-soluble drugs and for targeting specific sites in the body, current matrices do not exhibit the right balance of functional behaviour. For the new candidate drugs, polymers that have one set of properties in the oral cavity and another type of behaviour in designated areas of the GI tract may offer an answer to the delivery problem. Drugs of poor aqueous solubility can be formulated or conjugated to polymers that are soluble at high pH but hydrolyse at low pH to enable drug transport from the oral cavity to the stomach. The change in polymer properties can be in response to the local environment or induced by external stimuli such as charge, pH, light and temperature. Specific examples of these responses are considered later in the chapter. 3.2.2

Parenteral Drug Delivery

For parenteral drug delivery, for example intravenous (i.v) or intramuscular (i.m) injection, polymers can be used to enhance solubility, either directly or via the formation of higher order systems, such as micelles which can carry a water-insoluble drug molecule in their hydrophobic cores. Polymers can also be used for parenteral drug delivery in the form of micro and nanoparticles, as long as they meet the necessary criteria for injectable formulations of being sterilisable and stable against aggregation. The latter point is particularly important as material injected into the blood stream rapidly passes into the heart and then round the pulmonary circulation – any particles larger than  5 mm can lodge in pulmonary capillaries resulting in embolism. Thus, even for relatively simple polymers in parenteral use, careful consideration must be given to biopharmaceutical considerations and formulation requirements – for more complex responsive or actuating polymers the properties in the in vivo environment must be very tightly defined in order to prevent unintended adverse reactions. Nevertheless, the potential advantages of using ‘smart’ nanoparticles for parenteral delivery remain, as one can envisage a polymer surface coating that enables prolonged circulation in the bloodstream but which can be actuated to expose a binding moiety in response to a biological or external signal. Targeting in this case could be direct and instantaneous. Another facet of parenteral drug delivery is the implanted device, and this is perhaps the most promising and most readily commercialised area for responsive and/or active polymers. For an implanted vehicle or depot, drug release rate is controlled by dissolution and/or diffusion in the formulation, or for solid polymer implants by diffusion and/or degradation of the polymer. For more complex polymer hydrogels, the release can be controlled by the linking chemistries, and these can be made responsive to a wide variety of stimuli such as enzymatic action, redox potential and so on, as well as those noted above for the oral route. 3.2.3

Topical and Transdermal Drug Delivery

For other routes of delivery, the formulation considerations relating to polymers are less restrictive than for the parenteral route. Typically, for local rather than systemic activity, polymers can be used to enhance efficacy of topical drug delivery systems through solubility enhancement and or skin hydration. For transdermal delivery, polymers can have several roles, either as solubility modifiers, permeation enhancers or as aids in iontophoresis. The presence of multiple-charged species is a prerequisite for iontophoretic movement of drug molecules through the skin, so can be considered as an actuation process

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in its own right. However, clinically, iontophoresis is still a niche area in drug delivery and other routes are currently more attractive for development of responsive and actuator polymer systems, although some examples have been developed [18, 19]. 3.2.4

Delivery Challenges for Biomolecular Drugs and Cell Therapeutics

As noted earlier, the physicochemical properties of the drug and its biological target have important implications for the design of the correct delivery systems. The new classes of drugs emerging from the biotechnology sector offer challenges for delivery as their properties and formulation requirements, as well as their biological targets, can differ considerably from established small molecule counterparts. Irrespective of the design of the delivery device, there are barriers to delivery of these biomolecular drugs and cell therapeutics, and the most important of these are outlined below. 3.2.5

Peptides and Proteins

Peptides and proteins are inherently unstable due to the reactivity of amino acid side chains and a number of reactions can occur that can denature their primary, secondary and tertiary structures. These include transpeptidation (i.e. amide transfer reactions), hydrolysis of side chains from the backbone and racemisation. In redox environments generation of free thiols at cysteine residues can cause disulfide exchange, leading to changes in folds and tertiary structure, while uncoupling of the disulfides under reducing conditions can also lead to irreversible structural change if the protein is subsequently exposed to strong oxidants. This is because oxidation of thiol to sulfate can occur, in which case the presence of two sulfates arising from oxidised cysteines in close proximity will lead to charge– charge repulsion and possible chain rearrangement. Formulation of proteins within drug delivery devices and implants can lead to further loss in activity. For example, the introduction of a protein into a charged solid matrix, as might be required for an electroactive depot device, can result in denaturation of protein through irreversible adsorption into strongly charged regions. Preparation of nano or microparticle delivery systems that encapsulate proteins, within polymers such as poly(lactide-co-glycolide) (PLGA) or poly(caprolactone) (PCL), often involves the use of solvent precipitation or coacervation methods, and these can adversely affect proteins due to adsorption and precipitation at aqueous–organic interfaces [20]. Other factors, such as nonphysiological temperatures, pH values and ionic strengths, can also disrupt protein structure and hence subsequent function. It should be noted, too, that many polymer–drug delivery systems contain amphiphilic materials, for example PEO–PPO block copolymers (e.g. the commercially available Poloxamer and Pluronic classes), and these can exhibit strong detergent action, again leading to loss of protein activity. Manufacturing parameters are also important – for an implantable device or an injectable delivery system, the components must be purified and/or sterilised prior to use – with the obvious issue of thermal deactivation. Proteins are frequently required to be in a lyophilised form, again leading to irreversible aggregation and denaturation. In addition to the problem of adverse protein adsorption in the delivery matrix, many proteins adsorb to manufacturing vessels and plastic surfaces – thus careful design and control not only of the desired polymeric actuator used for drug delivery but also the processing and formulating environment to prepare it must take place.

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Nucleic Acids

The use of nucleic acids in therapy involves many of the hurdles experienced in protein delivery systems, including difficulties in formulating the delicate biopolymer drug and the high possibilities for degradation en route to the target. For DNA based therapies the target is the cell nucleus, whereas for RNA delivery to the cytoplasm is required. Many designed polymeric nucleic acid delivery systems (usually termed ‘vectors’) have been developed by analogy with viruses, which are of course natural DNA or RNA carriers that are able to deliver their payloads to cells with extreme efficacy. The requirements for these vectors are stringent; for example, for DNA, the nucleic acid (which can be several million Da in molar mass) must be condensed to a small size and protected from serum and intracellular nucleases. The DNA must be delivered to the target cell, cross the external cell membranes passively or actively, leave the endosomal compartments (avoiding degradative enzymes and escaping traffic to the lysosomes), then it must translocate into the nuclear compartment ready for transcription. The vector must thus protect the DNA and help transport it across multiple cell barriers, yet release it at the correct time. In addition, for an injectable DNA delivery formulation, the vectors with encapsulated/complexed DNA must be capable of extended circulation in the bloodstream in order to have chance to reach their cellular target. They must also be small enough to gain access to tissues and cells – typically this will put a size limit of <250 nm for the hydrodynamic diameter of the vector system in circulation. For commercial development, it is also desirable for a gene delivery vector to have flexible tropisms, so that one vector can be adapted for application across a wide range of disease targets. Above all, the vectors must be safe, that is nontoxic, nonimmunogenic and fully cleared from the body after use. It is the safety factor that has hampered the development of viral DNA and RNA delivery systems to date, as although viruses have evolved to be very successful in overcoming cell barriers, their therapeutic use is dependent on deactivation of their infective properties. Immune reactions are also a major problem that needs to be solved for longer term use of viral vectors in nucleic acid delivery. A fuller discussion of viral gene therapy is beyond the scope of this chapter, but the safety issues associated with viruses have prompted intense research into synthetic viral mimics. Transfection efficiency, acute and long-term toxicity and in vivo fate are all important challenges for synthetic polymer vectors as well as viral systems. Considering all the demanding requirements it is perhaps not too surprising that there are no successfully commercialised synthetic gene delivery vectors in the clinic so far. Later in this chapter some of the more specific challenges facing ‘smart’ polymer gene delivery vectors are described, but it should be noted that there are still a large number of these systems in the research and development stages, most likely as a result of the change in treatment paradigms that will occur should successful gene therapies reach the clinic. 3.2.7

Cell Delivery

Increasingly, cell and tissue/organ based therapies are being considered in medicine, as our knowledge of cell function and organisation develop, and as more challenging therapeutic targets are considered [21]. Tissue engineering and regenerative medicine strategies often involve the growth of cells and organ-precursors ex vivo on support materials. The aim of the process is ultimately to duplicate native tissues in terms of function, architecture and complex capillary networks.

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To culture new tissues, either natural or synthetic support structures, usually termed scaffolds, are used to provide anchorage for the cells and to impart mechanical stability and a degree of structural guidance. Ideally, the scaffold used should allow the cells to adhere, proliferate, differentiate and maintain their natural phenotype for the generation of a fully functional tissue or organ. If implanted, the scaffold materials used for cell culture must also promote integration of these artificial tissues as they grow with the surrounding tissue, by responding to physiological and biological changes. Such an ‘active’ scaffold can thus also be considered as a responsive cell delivery system and, as a consequence, stimuli-responsive or activated polymers are of interest as enhanced scaffold/support materials for regenerative medicine [22]. Biodegradable natural polymers such as alginate, hyaluronic acid and type I collagen have been used for applications that involve the implantation of the tissue complete with its scaffold. These form a temporary matrix around the cells, which gradually degrades in the body and leaves behind only the new tissue. However, studies in developing supports for cardiomyocytes have shown that scaffold degradation leads to the formation of fibrous tissue containing excessive amounts of extracellular matrix (ECM) and subsequent scar formation. This is similar to the damage seen in pathological disease states such as ischemic heart disease and dilated cardiomyopathy. Responsive or actuator synthetic polymers are potentially especially advantageous in this area, as their properties can be controlled to match the desired properties of the tissue they need to support; for cell types such as cardiomyocytes there is an obvious analogy in a repeated polymer response compared with the beating of cells in the heart. For all these seemingly contrasting types of drug delivery, including macromolecular and cell based therapeutics, there is a need to understand the relationship between the physiology of the body and the disease site, and the properties and functions of the materials used to deliver the therapeutics. It is also necessary to understand the distinction between controlled and site-specific delivery, especially the difference between release and targeting in order to design the optimum matrix or delivery vehicle. It should be noted that while control of drug release through a delivery system is desirable and can benefit site-specific accumulation, not all drugs are necessarily appropriate or suitable for targeting. Drugs that are not retained at a target site for a long enough period to act or which are rapidly degraded in situ will not benefit from site-specific release. In some cases, the cellular or organ target for therapy is also the site where toxicity is observed, thus site-specific release in these cases is not beneficial. In the case of certain biomolecular therapeutics such as antibodies, the specificity of interaction is such that additional targeting modalities are not required, but controlled or localised release can still be beneficial as breakdown of the antibody en route to the target can be reduced. For all drug delivery systems, a number of overriding criteria should be remembered. In order for a drug to exert its effect, it needs to be in physical contact with its physiological target (e.g. receptor), and therefore control in delivery is needed to ensure these interactions take place. This is also the case for cell delivery systems, although the definitions of physical contact are not on the same size (length) scale as with small molecule drugs. Adequate drug adsorption and adequate access to the target sites are important attributes of a delivery system, as are prevention of nonspecific drug distribution in the body, premature metabolism and premature excretion. In the context of ‘smart’ bioactive and electroactive polymers the ability to match drug input with required timing is perhaps the biggest advantage, and this opens up control of pharmacokinetics and biodistribution. As a consequence, many polymers are being developed that exhibit responsive or active behaviour.

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3.3 Emerging State-of-the-Art Mechanisms in Polymer Controlled Release Systems 3.3.1

Technologies for Controlled Drug Release

Polymers can be used to control release by mechanisms ranging from simple dissolution through to selectively triggered breakdown or externally activated conformational changes. The variations in release strategies reflect the different ways in which drugs, biomolecules and cells are associated with or attached to the delivery system. Delivery systems that dissolve to release a drug, or which rely primarily on matrix hydrolysis are not generally regarded as responsive or activated – and many other reviews of these systems are available. Thus the focus here is on systems where there is either a change in polymer conformation or a selectively cleaved triggered linkage as a working definition of an ‘active’ release system. 3.3.2

Polymer–Drug Conjugates

Polymer–drug conjugates are materials in which the therapeutic is covalently linked to a polymer backbone, and indeed for soluble polymers these are generally classed under the term ‘polymer therapeutics’ [2]. Polymer–drug conjugates have been primarily developed to date for cancer therapy as carriers for cytotoxic drugs, as a means to reduce side effects and expand dosage regimes [23, 24]. In particular, these conjugates exploit the high vessel permeability and poor lymphatic drainage exhibited by tumour cells [25, 26] which has been termed the Enhanced Permeation and Retention (EPR) effect [25, 27]. Many tumours exhibit increased but poorly formed vasculature, and the combination of this combined with poor efflux and leaky vessels results in the increased accumulation of polymers and particulates above a certain size range in the cancerous tissue. Although this is a passive effect, it can, nevertheless, serve to concentrate polymers in the tumour regions, and conjugation of low molecular weight cytotoxic drugs via a cleavable linkage to watersoluble polymers has already proved advantageous in early stage clinical trials [28–31]. Existing anticancer agents such as 5-fluorouracil (5-FU), platinates, doxorubicin, methotrexate, paclitaxel, podophyllotoxin and camptothecin have all been linked to soluble polymers to increase their local (i.e. target) toxicity while reducing adverse systemic effects [32]. Although polymer–drug conjugates have not generally been considered as responsive systems per se, pro-drug strategies and triggering chemistries accomplish a ‘switch’ in activity that can be locally activated, and thus very specific therapies can result. The conceptual basis of this approach has been demonstrated by studying the transport of polymer–drug conjugates into cells [33]. Polymer therapeutics have been shown to be strongly retained in multi-drug resistant cells, and the intracellular concentrations of drugs released from the polymers have been shown to be significantly higher after the same incubation time than when the free drug was administered [34]. 3.3.3

Polymer–Protein Conjugates

Attachment of polymers to peptides and proteins can greatly reduce instability of these biomolecules by resisting nonspecific proteolysis in vivo prior to arrival at the target site. Conjugation of peptides, most commonly to poly(ethylene glycol) (PEG), is a very effective

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way of ‘shielding’ the biomolecule from peptidases and the body’s immune system, without incurring immunogenicity or toxicity [35, 36]. Polymer–protein and polymer–peptide conjugates also exhibit prolonged circulation and reduced opsonisation in the bloodstream, and as a consequence a number of such conjugates are now being investigated or in clinical use. PEG has been widely used to produce new bioconjugated therapeutics and products in clinical use include PEG-L-asparaginase (Oncaspar), PEG-adenosine deaminase (Adagen), PEG-interferon (Pegintron and Pegays), PEG-filgastrim (Neulasta) and a PEGylated growth hormone (Pegvisomant). PEGylated anti-TNF-a has recently been approved for Crohn’s disease and continuing clinical trials are being conducted in a number of other therapeutic areas [37]. In addition to PEG based systems, polymers with responsive properties are being actively investigated for use in peptide and protein conjugates in order to obtain derivatives with enhanced biological functionalities or physicochemical and biopharmaceutical features [38–41]. 3.3.4

Polymer–Nucleic Acid Conjugates

Anti-sense and gene regulation based therapies requiring delivery of nucleic acids, siRNA and oligodeoxynucleotides (ODNs) suffer from many of the limitations that apply to protein delivery. Conjugation of polymers enhances the solubility of ODNs and can delay their degradation by nucleases, generally as a result of decreased access of nucleases to the polyphosphate backbone. Polymer–ODN conjugates have generally been prepared by phosphate ester formation at the 3’ or 5’ end of the ODN for PEGylated derivatives, [42] although more recent examples have incorporated specifically cleavable linkers between polymers and ODNs for site-specific delivery to subcellular compartments [43]. These results indicate the feasibility of this approach and suggest that, with more efficient targeting, polymer oligonucleotide conjugates should become useful anti-sense therapeutics. 3.3.5

Polymer–Nucleic Acid Complexes

As noted above, polymer–nucleic acid complexes are amongst the most widely investigated emerging drug delivery systems, and in many ways exemplify the challenges and opportunities in polymer based drug delivery devices. The classical ‘Trojan Horse’ analogy of drug delivery applies particularly well for responsive and triggered/activated drug delivery systems, as mechanisms must be sought for binding and hiding DNA/RNA during transport to the cell target, but expelling it once the site is reached. The permanent negative charge on each nucleotide unit in the nucleic acids allows the possibility for generating charge–charge complexes with polycations. The resulting polyelectrolyte complexes can be very stable against dissociation dependent on the number of associated charges, but of course any changes in charge, for example by redox or pH in the cell, or via an external stimulus, can correspondingly destabilise the interactions, leading immediately to a mechanism for switchable binding and release. In view of the potential importance of the polymer nucleic acid complexes in gene therapies and in the advantages that responsive polymers offer in this area, considered in depth here are the mechanisms by which the synthetic and biological macromolecules interact and how they must function in order to deliver and express a therapeutic transgene.

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Initial work in nonviral gene delivery vectors was carried out using cationic lipids and liposomes, and in many respects, both in terms of vector development and clinical trials, liposome mediated delivery has advanced to a greater extent than the cationic high polymer counterparts systems. However, polycation mediated delivery is advantageous over liposome formulations in other ways, in that liposomes not only interact with DNA but also with other lipid molecules. Hydrophobic segments in liposomes are key factors in determining the overall size, shape and stability, and the degree of interaction with DNA, cell membranes and with other lipids. There is, therefore, a limited ability to control liposomal structure without a functional penalty. Liposomes are also often poorly water soluble and tend to become unstable with time. Synthetic polymers by contrast have more easily controlled molecular architecture, and most are more easily modified with targeting ligands. Gao and Huang [44] demonstrated that the transfection efficiency of cationic liposomes could be enhanced by the high molecular weight polycation poly-L-lysine, by dramatically reducing the size of the complex, and possibly by offering better protection against nucleases, facilitating nuclear uptake and dissociation once inside the nucleus. Of course, gene delivery using polycations such as poly-L-lysine are successful vectors in their own right and many variants have been developed. 3.3.5.1

Polycation Mediated Delivery

Synthetic cationic polymers and their derivatives for gene delivery include synthetic peptides, poly-L-lysine (PLL), polyamines (such as polyethylenimine (PEI)) and polyamidoamine dendrimers, and poly(vinylimidazole) derivatives. These polycations self-assemble with DNA to form charge-neutralised or cationic complexes, dependent on charge ratio. Higher transfection efficiency and serum sensitivity is achieved with polycation systems compared to lipoplexes. Although gene transfer in some polycation vector systems, such as poly-L-lysine, is rather low compared to viral vectors, the use of the multiply-charged polyethylenimine has led to an effective gene delivery in many cell lines. To date, PEI remains the most effective polymer for gene delivery. It has been shown to target DNA in vivo and in vitro, to promote nuclear targeting and to facilitate DNA escape from endosomes. The action of PEI in part depends on its pH-responsive character, but PEI homopolymer is toxic at levels required for transfection, and thus other candidate polymers with pH or other responsive characteristics are required for medical use. An emerging area of research is the use of synthetic polymers that are able to bind DNA in a reversible manner employing functionality that bestows polymer conformation and hydration changes and DNA recognition. For nonviral systems to be effective vectors, the polymer must not only condense and deliver DNA intact to the target site but must also enable nucleic acid to be transported through the cell membrane, and be translocated from the cytoplasm to the nucleus. Many researchers have now investigated DNA complexed by electrostatic interactions to cationic polymers as a method by which the therapeutic gene can be transported. In most cases these cationic polymers form condensed complexes with DNA that both contract the nucleic acid to facilitate cellular uptake, and which protect it from serum and cytosolic nuclease degradation. Mechanisms of DNA condensation, cellular uptake and transport to the nucleus, as well as strategies to improve toxicity, transfection potential and nuclear targeting of polycation mediated delivery systems are discussed below.

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3.3.5.2

Mechanisms Of Nonviral Polycation Mediated DNA Delivery

DNA condensation is a naturally occurring process in vivo and is important for cellular processes, such as DNA replication and transcription. DNA condensation by polycations is of great importance if external DNA is to reach the nucleus of target cells. DNA is condensed by neutralisation of the negative charges along the phosphate backbone by the positively charged polycation. At a certain critical extent of charge neutralisation (around 90 % for polycations) DNA is condensed into small, tightly packed nanoparticles of between 20 and 200 nm diameter, through localised bending, resulting in the formation of a variety of condensed DNA structures, often reported to be toroid, rod-like or spheroidal. Indeed, it has been proposed by Pollard et al. that nuclear trafficking of complexed DNA is dependent more on the spherical morphology of condensed DNA, rather than on the ionic interactions [45]. This work is supported by Liu et al. [46–48] who observed that small, spherical complexes were more efficient for receptor mediated uptake. As well as localised changes in DNA structure, factors such as reduction in DNA segment interactions, as result of polycation binding, and DNA–solvent interactions contribute to its condensation into various morphologies. Several molecules of plasmid DNA may become incorporated into the condensed structure and thus it is often difficult to distinguish condensation from aggregation or precipitation. The size and morphology of condensed particles are dependent on many thermodynamic and kinetic factors. The formation of these particles is crucial for entry through the cell membrane, protection from nucleases and consequent transfection ability. The process of DNA condensation is reversible and can be defined as the dramatic decrease in volume (104-fold) occupied by a DNA molecule, usually with a finite size, and orderly morphology. There has been much experimental and theoretical investigation into the organisation of DNA into its condensed state, and the formation and arrangement of compacted DNA structures, including toroids, rods and intermediate structures [49–56]. As DNA condenses, it undergoes a coil–globule transition. Detailed fluorescence studies indicate that rods and toroids are formed as a result of the coil–globule transition of DNA. Once a toroid has reached a certain critical size, it can no longer package DNA in the usual way and the remaining coil condenses into an irregular globular structure. Many polymers are able to condense DNA to these toroidal and contracted nano/microsphere structures in vitro and the importance of efficient condensation for transfection has been demonstrated. There are a number of factors that influence DNA condensation and the size of polymer-DNA complexes. For example, low molecular weight linear PEI forms much larger complexes with DNA than higher molecular weight and branched PEI, and these complexes possess lower transfection abilities [57, 58]. 3.3.5.3

Cellular Uptake Of Nonviral Polymer–DNA Complexes

If the DNA–polymer complex carries an overall positive charge, it is able to interact with the negatively charged cell surface receptors of the plasma membrane and is taken up into the cell by endocytosis. After binding to cell surface receptors on the cell membrane, the complex is enveloped by invaginations and fusion of the cell membrane and this in turn form endosomes. The endosomes originate from areas of the cell membrane that are rich in two main types of specialised cell membrane components: coated pits and caveolae. Caveolae are thought to mediate DNA uptake to muscle, while coated pits are invaginations surrounded by membrane-associated proteins, including specialised cell-surface receptors,

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for specific cellular uptake (receptor mediated endocytosis). Once inside the membranebound endosome, any material (250 nm diameter) is usually processed by cellular mechanisms. Thus polymer–nucleic acid complexes must be strictly controlled in terms of their size to be therapeutically active. Endosomes become fused with golgi hydrolytic vesicles to form endolysosomes, and ingested material is broken down by hydrolytic enzymes activated by the acidity (pH 5–6) inside the active lysosome. As hydrogen ions are pumped into the lysosomes, an influx of chloride ions occurs, which relieves the accumulated proton gradient. The osmolarity within the lysosome increases, as a result of the influx of chloride ions. The resulting influx of water causes the lysosome to swell, and eventually burst, after degradation of ingested material. A simplified schematic of normal uptake by endosomes is outlined in Figure 3.1.

Figure 3.1 Schematic of endocytosis. Material that binds at cell surface receptors is ingested into the cell by invagination of the cell membrane, rich in the protein clathrin. An endosome forms, through sealing of the vesicle by fusogenic proteins. Once ingested material is taken up into the vesicle inside the cell, the surrounding protein is released. The endosome then fuses with a hydrolase vesicle derived from the golgi apparatus, to form an endolysosme. This causes activation of an ATP driven hydrogen pump that lowers the pH inside the vesicle, which in turn activates hydrolytic enzymes to digest enveloped material. Once the pH is lowered and hydrolytic enzymes are activated, the vesicle is termed a lysosome.

The most common uptake pathway into the cell for nontargeted polymers occurs via passive endocytosis, and indeed has been shown to be the pathway for both PEI and PLL [59, 60]. It is also likely that transfection occurring via endocytosis takes place for many polymers through calcium mediated cell anchorage to the extracellular matrix. The normal

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Figure 3.2 Schematic of cellular uptake of nonviral vector DNA complexes and evasion from lysosomal degradation; (a) shows the general proposed method and (b) illustrates more detailed mechanism of endolysosomal escape. In (b), the top figure outlines normal lysosomal degradation through active hydrolytic enzymes; due to the buffering capacity of many polycations, hydrolytic enzymes are not activated and the DNA–polymer complex is able to escape degradation (lower figure).

process of endocytosis and digestion by lysosomes, is potentially a barrier to nonviral delivery systems that rely on nonspecific cellular mechanisms to gain entry into the cell (Figure 3.2). A nonviral vector must be able to avoid degradation by lysosomes if it is to access the nucleus and be transcribed. Cationic polymers, such as PEI, are only partially protonated in the serum and cytosol but become fully protonated at low pH within the endolysosome as they accept and buffer the hydrogen ions that are pumped into the vesicle. The buffering action of PEI raises the pH within the endolysosome, resulting in inactivation of hydrolytic enzymes. In addition, the increased charge of protonated PEI results in an influx of chloride ions and water to counter the increase in osmotic pressure inside the endolysosomes and this can act to burst the membranes. It is likely that PEI and other polymers complexed to DNA are released from the lysosome as it ruptures before the hydrolytic enzymes can become activated and degrade the nucleic acid. The overall process by which PEI escapes through its protonation and endosomolysis is often termed the ‘proton sponge’ mechanism [61, 62]. Transfection of polymer–DNA complexes usually requires actively dividing cells and transfection efficiency is enhanced prior to mitosis, when the nuclear membrane has been broken down. Therefore, gene therapy with nonviral cationic vectors is advantageous in delivery to, for example, brain tumours, where normal cells do not divide and therefore are not transfected, but malignant cells undergo active cell division and are able to be targeted and destroyed. Brunner et al. demonstrated that linear PEI was not dependent on the cell cycle for transfection [63, 64]. 3.3.5.4

Nuclear Transport and Entry

After evasion from the endolysosome, DNA must traffic to the cell nucleus for transcription to take place. The mechanisms by which this is achieved in polymer–DNA complexes

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are complex and not fully understood. Suggested mechanisms include tracking through the cytoskeleton, while polymers that possess nuclear localisation sequences may be carried to the nucleus via importins. Work by Godbey et al. has shown that intact PEI–DNA complexes are able to enter the nucleus as discrete particles (possibly vesicular in structure) [65, 66]. Both PEI and PEI–DNA were shown to enter the nucleus, with a suggested mechanism involving coating of complexes with a lipophilic layer, which then fuses with the nuclear membrane. This coating may arise from the remains of the membrane-bound endolysosome or through electrostatic binding of anionic phospholipids to the cationic complex. Both coating mechanisms have been shown to take place with cationic liposomes and it is not unreasonable that nuclear uptake occurs in the same manner for other cationic polymers. For the design of polymer gene delivery vectors, it might be assumed that a mechanism must be encoded into the polymer to allow dissociation of the DNA from its polymer complex if it is to successfully be transcribed. However, some studies of PEI–DNA complexes have reported the complexes to be present in the nucleus in undissociated form, and thus unpackaging of the complexes must occur within the nucleus for these systems. Although the mechanisms remain unclear, DNA polymerase may be involved, and it has been postulated that, once inside the nucleus, polyanions such as endogenous DNA replace the transfected DNA within the complex in a polyion exchange reaction, releasing the therapeutic DNA within the nucleus. Synthetic polycations such as these may also behave in a way that mimics nuclear homing devices. Polycations may also be removed during transcription of nucleic acid by polymerases. Indeed, when injected directly into the nucleus, DNA complexed to PEI can be transcribed as efficiently as naked DNA. This was demonstrated by the work of Bieber et al. [67] who showed that DNA tightly complexed with PEI (at high N:P ratios) was transcribed DNA as well as naked DNA, and more efficiently as more loosely associated complexes at a low N:P ratios. This group therefore proposed that DNA does not need to be dissociated from its polycation before it is transcribed [67–69]. It can thus be concluded that nuclear localisation rather than DNA unpackaging in the nucleus is a key limiting factor in polymer mediated gene delivery and, as a consequence, design criteria for new polymer vectors, especially those that invoke responsive or activated mechanisms, need to take into account the need for nuclear targeting. Nuclear transport can be enhanced by modification of synthetic polymers with nuclear localisation signal peptides, [70–74] or with transcription factors, but of course this adds to the design and synthesis complexity of the vector system.

3.4 Responsive or ‘Smart’ Polymers in Drug Delivery 3.4.1

Soluble Smart Polymers

For the purposes of this section, ‘smart’ polymers are defined as those which exhibit a nonlinear response such as a conformational change or a phase transition to an external stimulus. A very large number of responsive polymers have been reported in both open and patent literature, but the field continues to grow as new mechanisms of response and new types of polymer are emerging.

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The prototypical ‘smart’ polymer is poly(N-isopropyl acrylamide) (P(NIPAM)), which exhibits an inverse temperature solubility profile in water, that is it is water-soluble below 32 °C but precipitates above 32 °C. The temperature at which this coil-to-globule phase transition occurs is known as the Lower Critical Solution Temperature (LCST), and conveniently this can be modified in P(NIPAM) by incorporation into the polymer chain of more hydrophobic or hydrophilic monomers. Owing to the fact that the LCST is close to body temperature and can readily be modified to just below or just above 37 °C through this co-monomer addition, P(NIPAM) polymers have been widely exploited in biomedical applications. The chemistries and applications of P(NIPAM) have been extensively reviewed elsewhere, [75–81] but even 15 years after one particularly well-cited review, many research groups are working with this remarkably versatile polymer [82–87]. Although P(NIPAM) materials have generated much interest in pre-clinical use, of particular note for the biomedical field are the recent reports from Lutz et al. on the synthesis and characterisation of a thermo responsive copolymer composed of oligo ethylene glycol units,

P(MEO2MA-co-OEGMA)

PNIPAm Cl

O

O O

x O O

*

y O

O

O

n* NH

O 9 O

Figure 3.3 Thermoresponsive polymers of P(MEO2MA-co-OEGMA) and P(NIPAM) (Reprinted ¨ . and Hoth, A. Point by Point Comparison of with permission from Lutz, J.-L., Akdemir, O Two Thermosensitive Polymers Exhibiting a Similar LCST: Is the Age of Poly(NIPAM) Over? J. Am. Chem. Soc., 128, 40, 13046–7. Copyright (2006) American Chemical Society).

poly(diethyleneglycolmethacrylate methyl ether-co-oligoethyleneglycolmethacrylate methyl ether) or P(MEO2MA-co-OEGMA) (Figure 3.3) [88]. This polymer is very promising for medical applications since it is made from biocompatible units and also because its LCST transition is sharper than that of P(NIPAM). There is no hysteresis observed for this polymer in repeated heat/cool cloud point cycles owing to the low interchain binding, whereas P(NIPAM) LCST behaviour is not fully reversible as a result of inter and intrachain hydrogen bonding across amide residues. Thus the P(MEO2MA-co-OEGMA) polymer has the capability to be even more precise in response to stimuli, as well as exhibiting enhanced biocompatibility [89–92]. In addition to temperature responses, polymer assembly and reorientation dependent on hydrophobic interactions and pH have been described. Block copolymer assembly into micellar structures is a potential means of encapsulating a therapeutic, and if disassembly into unimers can be triggered at a biological site, a powerful means for selective drug delivery can be envisaged.

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Kataoka and co-workers and the Eisenberg group pioneered the use of polymeric micelles in drug delivery applications [93–96]. The Kataoka group prepared copolymer micelles with the concept that these should exhibit properties similar to those of natural drug delivery systems – that is viruses. However, unlike viruses, multi-component micelles coated with polyethylene glycol are fully biocompatible and cannot be identified inside the body as foreign substances. It was found that therapeutic molecules could be inserted, such

flexible polymer brush biocompatibility and steric stabilization

driving force of core segregation hydrophobic interaction metal complexation electrostatic interaction

self-assembly in aqueous medium

amphiphilic block copolymer

several tens nm

reactive group introduction of targeting moiety

Figure 3.4 Idealised polymeric micellar containers for drug delivery (Reprinted with permission from Kataoka, K., Harada, A. and Nagasaki, Y. Block copolymer micelles for drug delivery: design, characterization and biological significance, Adv. Drug Del. Rev., 47, 1, 113–31. Copyright (2001) Elsevier Ltd).

as carcinostatic agents, metal complexes and even DNA, into the micellar core, suggesting that these systems could subsequently be used for drug delivery (Figure 3.4) [97, 98]. Control of micelle size is important in these systems, as it has been observed that 20 and 100 nm particles are effective in avoiding renal exclusion and reticuloendothelial uptake, and can be selectively targeted for tumours because of the high vascular permeability and EPR effect [99, 100]. It has been suggested that similar diblock copolymer assemblies may enter cells via an endocytosis process [101]. ‘Smart’ diblock copolymers that form differing types types of micelles in aqueous solution dependent on conditions of ionic strength and pH have been synthesised. These resulting two micellar states can be described as ‘schizophrenic’: by changing external pH, temperature or ionic strength the less hydrophilic block can transform to a hydrophobic state that forms the core of the micelle. By altering pH again, the second block becomes hydrophobic, effectively switching the micelles. Amphiphilic block copolymers of this type are of major interest for drug delivery as ‘dual triggered’ release systems, as it is possible to transform a container micelle into dissociated unimers to release a drug under a variety of conditions. The groups of Armes, Liu and Battaglia have been very active in this field and have reported a wide range of active or triggered polymer micelles [102–113]. An intriguing extension to the use of block copolymers involved triblock copolymers of PEG, and polyrotaxanes with b-cyclodextrins (b-CDs) [114, 115]. Varying stimulus

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response in these systems caused molecular contractions that resembled the action of myosin molecules sliding along actin filaments in the muscle contraction process. These triblock copolymers can be considered as intelligent biomaterials in that they can perform in similar ways to molecular structures in the body. For drug delivery it might be possible to protect a drug in one conformational state of the polyrotaxanes, but release the drug in another state. Other intelligent materials currently under development in the rotaxane area include tailored copolymers with end-functional groups and biodegradable sections that respond to specific biological conditions [116–120]. 3.4.2

Responsive Polymer–Drug Conjugates

Responsive polymer–drug conjugates have been prepared and adopted for active targeting in vivo; this can be achieved via a specific physiological response or external stimuli. The Urry and Chilkoti groups have used elastin-like polypeptides (ELPs), which exhibit very sharp phase transitions that can be tuned to a very high degree of specificity, for ‘smart’ behaviour [121–124]. Dreher et al. conjugated a thermo-responsive polypeptide to the anticancer agent doxorubicin via a pH cleavable linker. The conjugate was soluble in plasma at 37 °C but in hyperthermic conditions at 42 °C, as can be present in tumours or inducible via locally-directed ultrasound, the polypeptide exhibited a phase transition. The enhancement in lipophilicity increased the retention time of the conjugate. Therefore, by inducing hyperthermia at the site of action it was possible to increase the specificity of the treatment and decrease the adverse effects which an unbound drug would have in nontargeted tissues [125]. 3.4.3

Responsive Polymer–Protein Conjugates

Smart polymers have the possibility for optimising protein therapeutics through ‘hiding’ the protein during transit to cells, but ‘exposing’ or activating the protein as a result of a phase transition at the target site. Elegant work has established the concepts of ‘smart’ polymer–protein systems, [126–128] and recent studies have extended the concept to control of molecular motor proteins via responsive polymers [129, 130]. In general, the switching of enzyme activity is due to the change in steric demands at the active site imposed by the presence of a high molar mass polymer, with the chain extension and collapse controlling access or conformational restriction at the enzyme recognition site. However, the very factors that enable steric constraints to limit enzyme activity can be a problem in generating the conjugates in the first place. Thus, one major issue to overcome in these conjugates is the ability to attach a chain end of a highly mobile polymer to a particular site on the protein. Bontempo et al. [131] described a new approach to obviate this problem by growing P(NIPAM) from the protein streptavidin via a biotinylated polymerisation initiator (Figure 3.5). The key to successful polymerisation from the protein in solution was the use of sacrificial polymerisation initiators on an added insoluble resin. With the resin present the P(NIPAM) polymers grew in a controlled fashion from the protein although the polydispersity (PDI) was quite high compared to other polymers grown by atom-transfer radical polymerisation (ATRP). Nevertheless, this paper demonstrated the principle of growing polymers from biomolecule surfaces and this technique has much to offer for

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O HN

NH H N S

O

Br

O

O

O

Streptavidin macroinitiator

Streptavidin monomer, sacrificial initiator

Streptavidin macroinitiator

CuBr/bipy H2O

Streptavidinpolymer conjugate

Figure 3.5 PolymeriSation from streptavidin via a biotin initiator (Reprinted with permission from Bontempo, D. and Maynard, H. D. Streptavidin as a Macroinitiator for Polymerization: In Situ Protein-Polymer Conjugate Formation, J. Am. Chem. Soc., 127, 18, 6508–9. Copyright (2005) American Chemical Society).

protein drug delivery systems, as it should enable conjugates of very highly defined structure and functions to be generated. Further work by Heredia et al. involved activating bovine serum albumin (BSA) protein for ATRP via a disulfide linker. From this initiator, P(NIPAM) was grown in aqueous solution using a CuBr / bipyridyl system, both with and without a sacrificial initiator on a resin. The modification of the BSA was higher in presence of the sacrificial initiator, (65 % and 44 % respectively), while comparison of the polymers revealed that the PDI and molar mass were lower in the presence of resin [132]. The study also described P(NIPAM) growth from lysozyme. The growth was verified with size exclusion chromatography (SEC) and sodium dodecyl sulfate polyacrylamide gel electrophoresis (SDS–PAGE), although quantitative characterisation of the polymers wasn’t carried out in the latter part of the study. However, bioactivity tests indicated no loss in activity of the modified enzyme, showing that the growth of polymer–bioconjugates from proteins in situ is both feasible and promising [132]. In addition, the use of ATRP offers cleaner chemistry and easier purification of the material than conventional conjugation, factors that are likely to be crucial in take-up of polymer–protein conjugates for clinical applications. The recent introduction of other controlled polymerisation techniques to grow polymers from proteins [133–136] suggests that a host of new well-defined polymer–protein conjugates can be expected in the near future. 3.4.4

Responsive Polymers for DNA Delivery

The concept of using responsive polymer phase transitions to switch binding interactions with DNA was first explored by Hennink and co-workers, who showed that a cationic polymer of N,N’(2-dimethylamino)-ethylmethacrylate (DMAEMA) and related copolymers were able to bind and transfect DNA effectively. Copolymerisation of DMAEMA with NIPAm yielded polymers that complexed with DNA at room temperature but did not exhibit variations of binding with polymer LCST response. This group observed that low

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molecular weight copolymers, and copolymers with a high P(NIPAM) feed, aggregated at 37 °C and became unstable, leading to low transfection efficiency [137, 138]. However, complexes with higher molecular masses and with lower P(NIPAM) content were more stable at 37 °C and showed higher transfection efficiency. Thus, as the P(NIPAM) content of the copolymer increased, the copolymer:DNA ratio at which maximum transfection efficiency was reached, increased. Transfection of OVCAR-3 cells (human ovarian cancer cell line) using a b-galactosidase reporter gene showed that the number of cells transfected was about 10 % for polymer:plasmid weight:weight (w/w) ratios of 2–4. At weight ratios less than two, complexes were too large to be taken up by cells, and excess polymer at weight ratios greater than four induced toxicity into the cells. Low molecular weight copolymers were shown to be poor transfection agents, while high and ultra-high molecular weight copolymers showed the most effective transfection. Transfection efficiency also decreased with increasing P(NIPAM) content and zeta potential. Subsequently, Okano et al. demonstrated that transfection efficiency with related P(NIPAM) gene delivery systems was considerably increased by incorporation of a hydrophobic monomer unit into the responsive polymer backbone [139]. Incorporation of this hydrophobic unit (butylmethacrylate) also increased the weight ratio at which maximal transfection was reached. The Okano group additionally demonstrated that gene expression of b-galactosidase in COS-1 cells with DNA complexed with P(NIPAM)-co-DMAEMA-co-butylmethacrylate could be controlled by temperature, through modulation of complex association/dissociation. Further development of these polymers led to materials with enhanced DNA binding and improved transfection of DNA, with temperature dependent DNA binding demonstrated by gel retardation assays. By controlling the temperature at which transfection experiments were carried out, it was shown that control over gene expression could be achieved by careful polymer structure manipulation and transfection assay conditions [140, 141]. More recently, a number of groups have reported modulation of DNA binding and control of gene expression by responsive polymers, [142–151] and although the specific mechanisms and experimental conditions are somewhat varied, nevertheless, the principles of ‘smart’ synthetic gene delivery vectors have now been established.

3.5 Recent Highlights of Actuated Polymers for Drug Delivery Applications In recent years, a number of novel concepts have been developed in drug delivery wherein an electroactive response is used to trigger the drug release. This method has the advantage that it is essentially ‘bio-orthogonal’, that is that the stimulus is externally controllable and should not be affected by normal biological activity. Examples include applications in drug delivery and regenerative medicine [152–155]. Here two papers that best illustrate some of the powerful concepts in these novel stimulus-responsive polymer devices are highlighted. Wadhwa et al. prepared coatings from the conducting polymer poly(pyrrole) (PPy) that were designed to release an anti-inflammatory drug upon the action of an electrical

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Figure 3.6 Concept of poly(pyrrole) based charge-switchable drug delivery system.

stimulus [156]. The key to this concept is the switching of charge in PPy following cyclic voltammetry (CV) of the substrate (Figure 3.6). Electropolymerisation of pyrrole in the presence of an anionic drug, dexamethasone disodium phosphate (Dex), generated a positively charged PPyþ backbone with associated Dex– counterions. Addition of electrons to the system neutralised the polymer backbone, releasing the dexamethasone in a pulsatile fashion. In this paper, further control of release rate was obtained through the density of the films, with diffusion of the anion through the film being slow or via cyclic potential differences, which caused rapid swelling and deswelling of the film, acting as a rapid pump to release the drug. This group was thus able to obtain a linear correlation between the dose of the drug released from the film and the number of cycles of electrical potential. In vitro proof of efficacy was obtained using mouse glial cells, wherein inflammation pathways were controlled and no evidence of toxicity was obtained. George et al. have also explored PPy for drug delivery applications, but have introduced an ingenious method to enhance the range and types of drugs that can be released from a conducting substrate [157]. Poly(pyrrole) has been shown to be well tolerated by cells and, as described above, can bind anionic molecules following electropolymerisation. However, not all therapeutic drugs are anionic, and there is also a molar mass restriction on the types of molecules that can bind reversibly to a PPy surface or in a matrix of conducting polymer. The way round this problem was to use biotin as a linker between the PPy surface and the therapeutic of interest: biotin is a negatively charged relatively small molecule and is well known to exhibit very high binding to the proteins avidin and streptavidin. In addition, the presence of four binding sites for biotin at streptavidin enabled its use as a linker between PPy surface-bound biotin and biotinylated protein, in this case nerve growth factor (NGF). The overall concept is shown schematically in Figure 3.7. The key demonstration by George et al. was that biotin could be bound to the PPyþ substrate, then via linkage through streptavidin to biotinylated NGF the nerve growth factor protein was surface confined. Application of a voltage released NGF and the growth factor was shown to be active through an assay involving PC-12 cells, which are known to express a receptor for NGF that on binding leads to neurite extension.

80

Biomedical Applications of Electroactive Polymer Actuators (a)

–

+ Biotin Biotinylated NGF

PPy

Streptavidin

Platinum mesh

PBS

NGF concentration [ng/ml]

(b)

7

∗

Short Stimulation (30 seconds) Long Stimulation (150 seconds) No Stimulation No Biotin Dopant

6 5 4 3 2 1 0 Stimulation 1

5 min Stimulation 2 Incubation 1

5 min Incubation

Time point

Figure 3.7 Schematic of stimulated nerve growth factor (NGF) release. (a) Voltage applied across the polymer in PBS solution causes the release of the biotin from the PPy surface. (b) A 3 V stimulation of the PPy showed an increase in the amount of NGF released from the surface of the conductive polymer. Short stimulations of 30 s did not result in as much release as long stimulations of 150 s. The asterisk indicates a statistical difference versus short stimulation, no stimulation and no biotin dopant (p <0.05) (Reprinted with permission from George, P. M., LaVan, D. A., Burdick, J. A. et al. Electrically Controlled Drug Delivery from Biotin-Doped Conductive Polypyrrole, Advanced Materials, 18, 5, 577–81. Copyright (2006) Wiley-VCH Verlag GmbH).

3.6 Conclusions and Future Outlook In this short chapter it has not been possible to cover all the research areas underlying stimulus-responsive and active polymers for drug delivery, but the overall trends remain clear. The aim in all drug delivery applications is to take the therapeutic to the target and deliver it at the right time and in the right dose, thus improving the therapeutic index and reducing side effects, which in turn lead to better disease treatments. The dynamic and time dependent nature of this process, as well as the site-specificity requirement, all indicate that materials with some form of switchable or active properties are needed in order to adapt or respond to changing biological conditions. It is likely, therefore, that further developments in the synthesis, evaluation and applications of responsive polymers will continue in the drug delivery. Ultimately, the principal hurdles to the uptake of

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these materials in the clinic are the need for biocompatibility, the biodistribution and in vivo fate of the delivery systems, and the cost. However, with the ever-increasing burdens of complex diseases such as cancers in ageing populations, as well as emerging diseases in developing nations, the need for better drug delivery systems and more effective medicines will remain.

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4 Thermally Driven Hydrogel Actuator for Controllable Flow Rate Pump in Long-Term Drug Delivery Piero Chiarelli1 and Pietro Ragni2 1

Institute of Clinical Physiology, CNR, Italy Institute of Nuclear Chemistry, CNR, Italy

2

4.1 Introduction ‘Smart’ hydrogels can change their volume by a large fraction in response to many stimuli, such as temperature, pH, ions concentration, solvent composition and light irradiation. This environmental sensitivity plays an important role in achieving new technological and scientific applications. Hydrogels used as osmotic pumps were introduced in the 1970s. They allowed the development of new drug delivery systems, such as those used for the therapies of the gastrointestinal tract. Easy to regulate long-term devices with a very simple mechanical design could be commercially convenient with respect to those using complex electromechanical systems. In this chapter the use of a thermally activated hydrogel actuator for the realization of long-term drug delivery systems able to control the drug flow rate following defined tasks is investigated. In the first part, the material properties of the thermally sensitive hydrogel, whose network is constituted by cross-linked poly(vinyl-methyl-ether) (PVME) molecules, are outlined; the dynamic properties of a designed drug delivery system using such a hydrogel as the driving actuator are then derived.

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Biomedical Applications of Electroactive Polymer Actuators

4.2 Materials and Methods The hydrogel has been synthesized by means of free radical polymerization of PVME [1] by g–radiation. A PVME–sodium alginate emulsion was obtained by mixing a 30% (by weight) PVME aqueous solution with a 1% sodium alginate water solution. Then, the emulsion was cross-linked by means of Ca2þ ions in a 1 M calcium chloride (CaCl2) bath solution and submitted to g–radiation of 0.91 M rad/h for 24 hours. The final PVME hydrogel samples were cut in the form of parallelepiped of various dimensions and left to equilibrate in a bath of de-ionized water at room temperature. The sample’s equilibrium dimensions were measured by means of a stereo microscope as a function of temperature in a bath of de-ionized water, while the length–time behaviour of hydrogel samples submitted to free swelling experiments was detected by means of a Hall effect isotonic transducer connected to a computer acquisition data system.

4.3 Hydrogel Actuator The kinetic equation of motion of a hydrogel matrix is available in the frame of biphasic models [2, 3] and for simple cases it is possible to obtain the explicit time dependence of the spatially distributed strain of the gel [4–7, 9]. To obtain the relevant properties of the PVME actuator, it is necessary to know the material constants that define the kinetics of the hydrogel readjustment. In addition to the material constants, the hydrogel readjustment depends also on geometrical factors such as its shape and porous structure [8]. Our PVME actuator has a macroporous structure where the fluid is enclosed in pores, whose walls comprise a homogeneous gel given by the PVME polymer network and whose intermolecular spaces are filled by the water (interstitial fluid). The macroporous structure of our PVME gel makes the dynamics of the global actuator dependent by the physical length characterizing the pores and by their shape. In the PVME material made of connected cells of pseudo-spherical shape, the physical length is the mean thickness ‘a’ of the thin walls of the pores. From the equations of motion that describe the accommodation of thin gel layers [6, 9], the characteristic time (t) of the mechanical readjustment is proportional to the square of the wall thickness (‘a’) and inversely proportional to the gel diffusion coefficient (D), following the relation: D ¼ m=f

ð4:1Þ

where m is the shear elastic modulus of the gel and f is its friction coefficient, which in a neutral gel is given by the inverse of its hydraulic permeability [2, 7]. One of the simplest ways to determine the gel diffusion coefficient (D) is by means of free swelling experiment [5–7]. Since our actuator is driven by means of temperature changes, the free swelling kinetic is coupled to the heat transmission one due to thermal diffusion and interstitial fluid convection.

Thermally Driven Hydrogel Actuator for Controllable Flow Rate Pump

91

By observing that the time readjustment due to positive and negative temperature jumps does not change appreciably, it is possible to conclude that the mechanical response time of the PVME gel readjustment is much bigger than the one characterizing the thermal kinetics. The reason is that the thermal flow due to the interstitial fluid convection is of opposite sign in the gel shrinking with respect to the one in the swelling process. If the thermal kinetics were the limiting factor, a clear hysteresis would appear between the hydrogel contraction and expansion processes. By posing that the mechanical readjustment of the hydrogel is much slower than the thermal transmission kinetics, it possible to assume a quasi-instantaneous change of temperature of the whole actuator before starting the mechanical relaxation of the gel. 4.3.1

Thermo-Mechanical Gel Dynamics

To describe the spatially distributed strain of the PVME gel system, the THB equation of motion [5] can be used. Even if is an oversimplified version of the Biot’s poroelastic model [7], it adequately describes the diffusion kinetics of gel matrices [3, 5–7] and reads: f @ Ui =@t ¼ @ ij =@xj

ð4:2Þ

where Ui is the displacement vector of a gel element and f ¼ (K11 – K12 K21 / K22)–1 is the gel friction coefficient (where Kij are the electro-osmotic Onsager coefficients [7]); sij is the gel stress tensor that in the linear approximation reads: ij ¼ k eaa dij þ 2mðeij  eaa dij =3Þ þ adij

ð4:3Þ

where k and m are, respectively, the bulk and the shear elastic moduli of the gel, a is the chemically or thermally induced stress at zero strain for isotropic materials and dij is the Kroneker delta; eaa is the gel dilatation, given by the trace of the strain matrix eij: eij ¼ ð@Ui =@xj þ @Uj =@xi Þ=2 In the case of an uncharged polymer network (i.e. K12 ¼ K21 ¼ 0), the friction coefficient results: f ¼ (K11)–1 [7], where K11 is the gel hydraulic permeability. Generally speaking, the material parameters in Equation (4.1) and Equation (4.2) are functions of the physical variables of the material (e.g. temperature, strain itself, etc.) as well as of the chemical ones (e.g. pH, ionic strength and type of solvent). Because the dependence of the material parameters on the mechanical deformation is weak, m, k and f can be assumed constants in the thermal gel readjustment, where the temperature is held constant for t > 0. Moreover, since the temperature change can be approximately assumed uniform over the sample, the thermal stress (a) is only a function of time. Given a(T, t £ 0) ¼ a0, the application of a sudden change of temperature (DT) at time t ¼ 0 that is held constant as a function of time, leads to the following thermal stress function: aðt Þ ¼ a0 þ yðt Þ ðaðT þ DT Þ  a0 Þ By choosing the null reference thermal stress state at t > 0 so that: aðt > 0 Þ ¼ aðT þ DT Þ ¼ 0

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Biomedical Applications of Electroactive Polymer Actuators

the following equality is obtained: aðT; t Þ ¼ ð1  yðt ÞÞa0 where y(t) is the unit step function. Hence, the application of a temperature step change is equivalent to freeing the gel sample, at time equal zero by the stress a0 , and letting it move toward the equilibrium volume, such as in a free swelling experiment. In the case of thermal stimulation, it is possible to perform free deswelling (a0 < 0) experiments depending on the sign of temperature change. By introducing Equation (4.2) in Equation (4.1), and by taking the divergence of both members and inserting the incompressibility of solid and liquid constituents, the equation of motion finally reads [7]: f @ eaa =@t ¼ ðk þ 4m=3Þ@ 2 eaa =@xi @ xi

ð4:4Þ

When the gel sample has the shape of a thin quasi-planar layer, assuming the z-axis is perpendicular to the gel layer plane, Equation (4.2) can be simplified as follows [6, 9]: @exx =@t ¼ D @ 2 exx =@z2

ð4:5aÞ

@eyy =@t ¼ D @ 2 eyy =@z2

ð4:5bÞ

@eaa =@t ¼ Db @ 2 eaa =@z2

ð4:5cÞ

where D ¼ m/f and Db ¼ (k þ 4m/3)/f. The spatio-temporal solutions for the strains exx and ezz are [9]:

exx ¼

     1  4e0 X ð1Þn ð2n þ 1 Þ 2 t ð2n þ 1Þzp cos exp t a p n ¼ 1 2n þ 1

ezz ¼ eaa  2exx ¼

   1  4e0 X ð 1Þ n ð2n þ 1Þzp cos a p n ¼ 1 2n þ 1 

3 exp



ð4:6aÞ

ð4:6bÞ

   ð2n þ 1 Þ 2 t ð2n þ 1 Þ 2 t 2 exp tb t

where t ¼ a2 / p2D and tb ¼ a2 / p2Db are the characteristic time constants for the ‘shear’ and ‘bulk’ diffusional gel readjustment, e0 is the initial uniform strain of the sample with respect the final one (at t ¼ 1) assumed as reference (e1 ¼ 0) and ‘a’ is the gel layer thickness at t ¼ 1. Moreover, since Db > D, it follows that tb < t From Equations (4.6a) and (4.6b), the length L(t) and the thickness a(t) of the gel are obtained as a function of time, respectively, to read [9]:

Thermally Driven Hydrogel Actuator for Controllable Flow Rate Pump

Lðt Þ ¼ L1

93

(  )  1   4e0 X ð  1Þ n ð2n þ 1Þ 2 t 1 þ exx ð Z ¼ 0 Þ ¼ L1 1 þ exp t p n ¼ 1 2n þ 1

(

aðt Þ ¼ a 1 þ 2

Z

a=2

ezz 0

ð4:7aÞ

)

(

   )  1  8e0 X ð 1 Þ n ð2n þ 1 Þ 2 t ð2n þ 1Þ 2 t ¼a 1 þ 2 2 exp 3 exp tb t p n ¼ 1 2n þ 1

ð4:7bÞ

Given the above kinetics for the pore walls and assuming the global macroscopic gel dimension quasi-isomorphic to the pore diameter, and hence to the pore wall circumference, the gel actuator length will approximately follow the time law given by Equation (4.7a). For t > 9t in Equation (4.7a) the slower exponential relaxation prevails, so that the PVME actuator length reads:   2   h t i 4e0 4e0 p D Lðt Þ ffi L1 1  exp exp 2 t ð4:8Þ ¼ L1 1  t a p p By fitting the exponential length relaxation of the PVME samples it possible to obtain the characteristic time (t) and the gel diffusion coefficient (D). 4.3.2

Experimental Results

The dependence of the PVME response time constant, t ¼ a2/p2D, as a function of the temperature is reported in Figure 4.1. By introducing the temperature dependence of the

Response time constant (s)

20

• •

15

•

•

•

•

10

•

5

• •

0 20

25

30

35

40

45

Temperature (°C)

Figure 4.1 Readjustment time constant of the PVME gel as a function of temperature.

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Biomedical Applications of Electroactive Polymer Actuators

pore wall thickness, a ¼ a f(T), where f(T) ¼ L(T)/L(T ¼ 20°C) is reported in Figure 4.2 and where a ¼ a(T ¼ 20°C), the normalized diffusion coefficient, D/a2, shown in Figure 4.3 is obtained. The data refer to thermal free swelling experiments at different temperatures by applying a thermal sudden jump of 2 °C.

•

Fractional PVME length

1.0

•

• •

•

•

•

•

•

0.9

•

• • • • • •

0.8

• • •

0.7

• • •

0.6 15

20

25

30

• •

35

• •

40

45

Temperature (°C)

Normalized diffusion coefficient (s–1)

Figure 4.2 The equilibrium length L(T) of the PVME gel normalized by one at T ¼ 20 °C as a function of temperature.

0.014

•

0.012 0.010 •

0.008 0.006

• •

• •

0.004

• •

•

0.002 20

25

30

35

40

45

Temperature (°C)

Figure 4.3 The shear diffusion coefficient D/a2 of the PVME gel matrix as a function of temperature.

Thermally Driven Hydrogel Actuator for Controllable Flow Rate Pump

95

By using the relationship that links the Young’s elastic modulus (E) to the bulk (k) and to the shear (m) elastic moduli that reads: m ¼ Eð3=ð9  E=kÞÞ since, in gels, it is usual for k > E [1, 6, 8], it follows that: 3=8

E < m < 1=3 E

Therefore, by measuring the Young’s elastic modulus (E) of the PVME by means of independent force–elongation experiments, reported in Figure 4.4, it is possible to evaluate the shear elastic modulus of the material, reported in Figure 4.5, with a precision of about 5 %. Once the shear elastic modulus has been determined, as reported in Figure 4.6 the friction coefficient ‘f ’ is obtained by the relation D ¼ m /f. It is interesting to note that the shear elastic modulus (m) and the friction coefficient (f) show the typical dispersion sigmoid and bell shape, respectively.

Young’s elastic modulus (N/s2)

60000

• • •

50000

40000

•

30000 •

20000

•

• •

•

10000 20

25

30

35

40

45

Temperature (°C)

Figure 4.4 The Young’s elastic modulus (E) of the PVME gel as a function of temperature.

The force generated by a PVME gel strip of a known sectional area, is a consequence of the thermal stress (a(T)) of Equation (4.3). Actually, the effect of the thermal stress (a(T)) is to change the free gel rest length and, therefore, it is possible to obtain the isometric force generation of a PVME actuator, with free lateral surfaces, by means of the equilibrium PVME length measured as a function of the temperature shown in Figure 4.2 with the Young’s gel modulus reported in Figure 4.4. By posing that the PVME actuator exerts zero force at 36 °C, the stress generated by its temperature change is reported in Figure 4.7.

96

Biomedical Applications of Electroactive Polymer Actuators

•

Shear elastic modulus (N/s2)

20000

• •

15000 •

10000

• • • •

•

5000

20

25

30

35

40

45

Temperature (°C)

Figure 4.5 The shear elastic modulus (m) of the PVME gel as a function of temperature.

Friction coefficient (Ns/m2)

3.5 × 106

•

3.0 × 106

•

• •

2.5 × 106

•

2.0 × 106 1.5 × 106

• •

1.0 × 106

•

20

•

25

30

35

40

45

Temperature (°C)

Figure 4.6 The fluid–matrix friction coefficient (f) of the PVME gel as a function of temperature.

Thermally Driven Hydrogel Actuator for Controllable Flow Rate Pump

97

PVME stress generation (N/m2)

10000 •

5000

• • •

•

0 •

–5000 • •

–10000 20

25

30 35 Temperature (°C)

40

•

45

Figure 4.7 The generated stress of the PVME gel actuator as a function of temperature when its zero force point is chosen at T ¼ 36°C.

4.4

Pump Functioning

The schematic drawing of a controllable pump for long-term drug delivery is shown in Figure 4.8. An internal spring is regulated for the maximal drug release that is achieved at the lowest device temperature (body temperature of 36 °C). Then, to lower the drug flow as requested, the temperature of the actuator must be appropriately raised by means of the thermal heating of the resistors. Since cooling of the PVME cell happens spontaneously and the lost heat power is fixed, while the heating power can be regulated by the electrical energy dissipated into the resistors, the heating–cooling cycle is not symmetric. Therefore, if the PVME contraction can be practically as fast as desired, its relaxation time is fixed and slow. This fact limits the application of the PVME gel motor (as all thermally driven mechanisms such as shape memory alloys) to phenomena that do not have very fast kinetics. The use of Peltier’s cells that actively cool the gel mover can lead to a faster relaxation response of the actuator and increase of drug delivery. An internal programming unit able to take into account of the thermal and mechanical inertia of the whole pumping system will definitely improve the timely regulated drug outflow.

98

Biomedical Applications of Electroactive Polymer Actuators long term reservoir

36°C

long term reservoir

PVME actuator

38°C

heaters

passive spring

long term reservoir

42°C

Figure 4.8

Schematic drawing of a controllable pump for long-term drug delivery.

4.5 Conclusion In this present chapter it has been shown how a thermally controlled hydrogel actuator can be used to let a drug infusion pump execute a time-defined tasks. By using a biphasic model, the characteristics of the hydrogel mover (as force density and time response) are explicitly defined together with their functional dependence by the geometrical and material parameters. The macroporous structure of the ‘hydrogel motor’ allows a quick contractile response to temperature changes regardless of its dimension. The use of thermoelectric cooling units can shorten the thermal cooling and the elongation time of the actuator.

References 1. Hirasa, O., Morishita, Y., Onomura, R., et al. (1989) Preparation and mechanical properties of thermo-responsive fibrous hydrogels made from poly(vinyl methyl ether)s, Kobunshi Ronbunshu, 46, 661–5.

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99

2. Biot, M. A. (1956) Theory of propagation of elastic waves in a fluid-saturated porous solid, I low-frequency range, J. Acoust. Soc. Am., 28-2, 168–90. 3. Johnson, D. L. (1982) Elastodynamics of gels, J. Chem. Phys, 77, 1531–9. 4. Peters, A. and Candau, S. J. (1988) Kinetics of swelling of spherical and cylindrical gels, Macromolecules, 21, 2278–82. 5. Tanaka, T. and Fillmore, D. J. (1979) Kinetics of swelling of gels, J. Chem. Phys, 70, 1214–8. 6. Chiarelli, P. and De Rossi, D. (1988) Determination of mechanical parameters related to the kinetics of swelling in an electrically activated contractile gel, Prog. Coll. Polym. Sci., 78, 4–8. 7. Chiarelli, P. and De Rossi, D. (1992) Modeling and Mechanical Characterization of Thin Fibers of Contractile Polymer Hydrogel, J. Intelligent Materials System and Structures, 3, 398–417. 8. Suzuky, M. and Hirasa, O. (1993) An Approach to artificial Muscle Using Polymer gels Formed by micro-phase Separation, Adv.Polym. Sci., 110, 241–61. 9. Chiarelli, P., Domenici, C. and Genuini, G. (1993) Crazing Dynamics in the swelling of thermally cross-linked PVA–PAA films, J. Mat. Sci.: Mat. in Med., 4, 5–11.

Section II Ionic Polymer–Metal Composites (IPMC)

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

5 IPMC Actuators: Fundamentals Kinji Asaka and Keisuke Oguro National Institute of Advanced Industrial Science and Technology (AIST), Japan

5.1 Introduction Ionic polymer–metal composites (IPMCs), which are composed of ionic gel polymer plated with metal electrodes, are one of the most promising electroactive polymer (EAP) materials for the artificial muscle-like actuators. The image and the schematic drawing of the actuation of the IPMC driven by low voltages are shown in Figure 5.1. When applying a voltage, the counter cation moves the cathode side with dragging water, which results in the pressure gradient for the ionic gel polymer. A large ionic current gives the IPMC actuator a soft and relatively powerful motion, as shown in Figure 5.1. IPMC actuators have number of advantages that make them attractive to use for various biomedical and human affinity applications:

     

Low drive voltage (1–3 V) Relatively high response (up to several hundreds of Hertz) Large response Soft material The possibility and ease to miniaturize and to mould into any shape Can be activated in water or in wet condition. Possible to work in dry condition.

Historically, direct transformation from electrical energy to mechanical work using ionic polymer gel was firstly reported by Hamlen et al. in 1965 [1]. After that, many pioneer workers investigated the electric response of ionic gels [2]. Following these workers, Oguro et al. [3] firstly reported the bending response of the perfluorosulfonic acid membrane (Nafion 117) plated with platinum electrodes, being activated by low

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Biomedical Applications of Electroactive Polymer Actuators

Weight

Voltage IPMC actuator 1second

Electrode

Voltage

Ionic conductive Polymer : Polymer electrolyte (Anion)

: Water molecule

(a)

: Cation

(b)

Figure 5.1 (a) Photograph of the bending performance of the IPMC actuator. The Nafion/Au composite of 0.8 mm thickness lifts 10 g weight driven by a 3 V voltage. (b) Schematic representation of the structure of the IPMC actuator.

voltages, about 1 V, in 1992, which was named IPMC [4]. Shahinpoor et al. also reported a similar idea in 1992 [5]. The IPMC actuator was more durable and it had higher response than the electric response polymer gels that were known at that time. Hence, many researchers have been applying IPMC actuators to various applications since then. The IPMC is also known as the ionic conductive polymer gel film (ICPF) [6]. Described in this chapter, are the basic aspects of IPMC actuator fabrications, measurement methods for testing, actuator performances, physics-based models and recent development of the material of the IPMC-like ionic polymer based actuators, based on our previous works. Interested readers can refer more comprehensive review articles on the IPMCs [6–9].

5.2 Fabrication 5.2.1

Ionic Polymer

Ionic polymers usually used for the IPMC are perfluorosulfonic acid or perfulorocarboxylic acid polymers, of which the typical chemical structures are shown in Figure 5.2 [10]. Commercially available products of thin films made from perfluorosulfonic acid can be obtained from E.I. Dupont de Nemours Co. (Nafion). Several other companies supply similar compounds. Asahi Glass Co. produces perfluorocarboxylic acid type (Flemion). A thin film of the perfluorinated ionic polymers can be obtained by casting their dispersing solution and evaporating the solvent, or by hot moulding a thermoplastic form of their polymers and changing –SO2F to –SO3– by hydrolysis. The device can be fabricated in any shape by using the casting solution or the thermoplastic beads of commercially available products of perfluorosulfonic acid (Nafion) (Chapter 6).

IPMC Actuators: Fundamentals –(CF2

CF2)n –(CF

105

CF2)–

O CF2 CF–CF3 O (CF2)y X

Figure 5.2 Chemical structure of prefluorinated ion exchange polymer. X ¼ SO3–: perfluorosulfonic acid polymer; COO – perfluorocarboxylic acid polymer; or SO2F: thermoplastic polymer.

Hydration of the fluorinated ion exchange resins depends on the ionic form and ion exchange density [11–13]. Figure 5.3 shows the water content of the perfluorosulfonic acid (Nafion 117 (N-117), charge density 0.91 meq./g) and perfluorocarboxylic acid membranes (Flemion (F-1.44), charge density 1.44 meq./g, and Flemion (F-1.8), charge density 1.8 meq./g) of various ionic forms. In the case of the Nafion membrane, the water content decreases as the hydrophobicity of the counter cations increases. F-1.44 has lower water content than N-117, while F-1.8 has much higher water content in the case of every counter cation. The ionic conductivity of the three kinds of membranes, which were estimated by impedance measurement, are shown in Figure 5.4. The ionic conductivity also depends on the ionic size, the hydration, the charge density of the ionic polymer, and so on. The alkali and alkali earth cation-form polymers have larger conductivity than the alkyl ammonium cation-form polymers. As the size of the alkyl ammonium cation increases, the conductivity decreases. Flemion membranes have larger conductivity than Nafion membranes. These properties, in relation to the electric bending response of IPMC actuators, are discussed in detail later. 5.2.2

Plating Methods

The plating electrodes for optimum performance of the IPMC actuator should have the following criteria:

      

Good adhesion to the ionic polymer High electric conductivity Large electrochemical interfacial area Large electrochemical window (high over potential to redox reactions) Inertness Softness Nontoxicity.

An established method of electrode plating on the ion exchange membrane for fulfilling the above criteria is chemical plating with platinum or gold electrodes. Oguro et al. firstly found the bending response of Nafion 117 chemically plated with platinum electrodes. However, a platinum electrode has a narrower electrochemical window and is mechanically

Biomedical Applications of Electroactive Polymer Actuators g water/g dry membrane (%)

106

35 30 25 20 15 10 5 0

Li

a

N

K

s

C

C

a

Ba

g

M

A

TE

A

A rA TE TP

A TB

A

A rA TE TP

A TB

TM

A

r TP

A

A

TB

water content g water/g dry membrane (%)

(a) 30 25 20 15 10 5 0

Li

a

N

K

C

s

C

a

g

M

Ba

TM

g water/g dry membrane (%)

(b)

60 50 40 30 20 10 0

Li

a

N

K

C

s

C

a

g

M

Ba

TM Membrane Form (c)

Figure 5.3 Water content of (a) Nafion 117, (b) Flemion F-1.44 and (c) Flemion F-1.8 membranes of various ionic forms. Abbreviations used for alkyl ammonium ions are TMA, TEA, TPrA, TBA for tetramethyl, tetraethyl, tetrapropyl and tetrabutyl ammonium ions, respectively (Reprinted with permission from Asaka, K., Fujiwara, N., Oguro, K. et al. State of water and ionic conductivity of solid polymer electrolyte membranes in relation to polymer actuators, J. Electroanalytical Chem., 505 (1–2), 24–32. Copyright (2001) Elsevier).

harder than a gold electrode. Hence, a chemical plating method with a gold electrode has been developed for the IPMC actuator [14]. Two different methods, known as ‘reductant permeation’ (RP) and ‘impregnation reduction’ (IR), have been successfully developed for plating the electrodes for the ion exchange

IPMC Actuators: Fundamentals 0.1

K N-117

Na

Li

0.001 TEA

Mg Ca Ba

membrane conductivity (λ m /S cm–1)

membrane conductivity (λ m /S cm–1)

0.01

Cs

TMA

0.0001 –5

10

10–6

107

TPrA

TBA

F-1.44 NH4

0.0001

TPrA

TEA

Cs

Ca

TMA

0.001

K

Na

Li

0.01

Mg Ba

TBA

10–7 10 20 30 40 50 60 70 80 limiting equivalent conductivity (λ w /Scm2equiv–1)

10–5 10 20 30 40 50 60 70 80 limiting equivalent conductivity (λ w /Scm2equiv–1)

(a)

(b)

membrane conductivity (λ m/S cm–1)

0.1 F-1.8 K

Na

0.01

Li

Ca

TMA

TPrA

Cs

Mg

TEA

Ba

0.001 TBA

0.0001 10 20 30 40 50 60 70 80 limiting equivalent conductivity (λ w /Scm2equiv–1)

(c)

Figure 5.4 Dependence of the membrane conductivity of Nafion 117 (a), F-1.4(b), and F-1.8 (c) membranes of various ionic forms on the limiting ionic conductivity of each ion in water. Abbreviations used for alkyl ammonium ions in the figure is the same as Figure 5.3 (Reprinted with permission from Asaka, K., Fujiwara, N., Oguro, K. et al. State of water and ionic conductivity of solid polymer electrolyte membranes in relation to polymer actuators, J. Electroanalytical Chem., 505 (1–2), 24–32. Copyright (2001) Elsevier).

resins under wet conditions. In the RP method, a metal layer is formed on the surface of a membrane by permeation of reducing agents into the other side of the membrane when a metal complex solution and a reducing solution are placed on either side of the membrane, respectively. In the IR method, the cation exchange membrane with pre-exchanged cationic metal species is subsequently immersed in the reducing solution, which reduces and displaces the metal toward the outer surfaces of the membrane. The IR method is known to be better than the RP method for achieving the above criteria. Fujiwara et al. [14] firstly developed the gold plating method on the surfaces of the ion exchange membrane by using the IR method. The schematic representation of the chemical plating of gold on the surfaces of the ion exchange resin by the IR method is shown in Figure 5.5. After roughening the surface of the ion exchange resin membrane by dry blasting or emery paper, the ion exchange polymer is

108

Biomedical Applications of Electroactive Polymer Actuators Ion-exchange polymer

Au [AuL]+ SO3–

SO3– H+ +[AuL]+

H+ SO3–

H+

–H+

SO3–

SO3–

SO3–

[AuL]+

+X+

[AuL]+

–L

SO3–

SO3–

1.Ion-exchange

SO3–

X+ X+ SO3–

SO3–

X+

[AuL]+

H+

Au X+

2. Reduction

SO3– 3. Sequential

+ [AuL]+ : N

N Au Cl

Cl

Figure 5.5 Chemical plating of gold onto the surface of the ion exchange polymer using a cationic gold complex such as dichlorophenanthrolinegold (III) (Reprinted with permission from Fujiwara, N., Asaka, K., Nishimura, Y. et al. Preparation of Gold-Solid Polymer Electrolyte Composites As Electric Stimuli-Responsive Materials, Chem. of Mat., 12 (6), 1750–4. Copyright (1999) American Chemical Society).

immersed in the aqueous solution of gold complex salt. Then, the metal complex adsorbed in the ion exchange polymer is reduced by the reducing agent. The sequential plating technique has been developed for optimizing the electrode structure, which has large electrochemical area and soft mechanical property.

5.3 Measurement The measuring setup for the displacement of the IPMC actuator in response to electric signals is shown in Figure 5.6(a). Typically, the actuator strip – 15 mm length and 1 mm width – is clamped by two gold disks and the displacement at a point 10 mm away from the fixed point (free length) is measured by the laser displacement meter. The applied voltage and electric current are simultaneously measured. The curvature is evaluated from the measured displacement () by the following equation: 1=R ¼ 2=ðl2 þ 2 Þ;

ð5:1Þ

where R is the curvature radius and l is the free length. The blocking force is also measured by the setup, as shown in Figure 5.6(b). The relationship between the bending force (F) and the curvature (1/R) is: 1=R ¼ ðM

FlÞ=EI

ð5:2Þ

IPMC Actuators: Fundamentals

109

Load cell

Waveform generator/ potentiostat

Waveform generator/ potentiostat Electrode

Electrode Actuator element

Curvature radius: R Laser displacement meter

Laser

Actuator element

l

Force

Displacement: δ

(a)

(b)

Voltage (V)

1 0 –1

Current (mA)

20 0 –20

Displacement (mm)

Displacement (mm)

Current (mA)

Voltage (V)

Figure 5.6 Schematic representations of the measuring setup for the performance of the IPMC actuator. (a) Displacement measurement. (b) Force measurement.

0.2 0.0 –0.2 0

10

20

30

40

50

60

70

80

90

1 0 –1 20 0 –20

0.2 0.0 –0.2 0

10

20

30

40

50

60

Time (s)

Time (s)

+ a. Na form

b. (C4Hg)4N+ form

70

80

90

Figure 5.7 Performance of the Flemion/Au actuator: (a) Naþ form; (b) (C4H9)4Nþ form (Reprinted with permission from Onishi, K., Sewa, S., Asaka, K. et al. The effects of counter ions on characterization and performance of a solid polymer electrolyte actuator, Electrochimica Acta, 46 (8), 1233–41. Copyright (2001) Elsevier).

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where M is the bending moment that is a function of input voltage, EI is the rigidity of the IPMC. In the case of the displacement measurement, Equation (5.2) is: 1=R ¼ M=EI

ð5:3Þ

since F ¼ 0. In the case of the blocking force measurement, Equation(5.2) is: F ¼ M=l

ð5:4Þ

since 1/R ¼ 0. The bending moment as a function of input voltage can be derived by the microelectromechanical transfer model, which is described in Section 5.5.

5.4 Performance of the IPMC Actuator During the past decade, much work has been done on the development of the IPMC actuator. In this section, typical examples of our previous experimental results with the Nafion/Au and Flemion/Au actuators are summarized. Shown in Figure 5.7 are the experimental curves of the input voltage, the electric current and the displacement of the Flemion/Au composite of the sodium ion (Naþ) form (a) and the tetrabutyl ammonium ion (C4H9)4Nþ from (b) [15]. The displacement response of the Nafion 117/Au composite actuators of the lithium (Liþ), Sodium (Naþ), Cesium (Csþ) and tetraethyl ammonium (TEA) ion (C2H5)4Nþ forms are shown in Figure 5.8 [16]. The displacement curves of alkali ion form of both Nafion and Flemion are a typical response of the IPMC actuator in application of step voltages. When applying the step voltages, the actuator strip bends to the anode side quickly (within 0.1 s), then slowly back to the cathode side (back motion), and finally stops in a stable position. When using an anion exchange membrane as the ionic polymer, the reverse behaviour takes place [17]. The back motion depends on the counter cation species of the Nafion or Flemion membranes as shown in Figures 5.7 and 5.8. When bulky alkyl ammonium cations, such as tetraethyl ammonium or tetrabutyl ammonium, are used, there can be no back motion as shown in Figures 5.7b and 5.8. The ionic forms of the Nafion and Flemion affect the response speed and the bending amplitude of the electromechanical response of the IPMC actuator as shown in

Displacement (mm)

8 6 4 2 0 0 –2 –4 –6 –8

Li Na Cs

TEA 0.1

0.2

0.3

0.4

0.5

Time (s)

Figure 5.8 Displacement response of the Nafion/Au actuator with Liþ (Li), Naþ (Na), Csþ (Cs) and (C2H5)4Nþ (TEA) as counter ions [16].

IPMC Actuators: Fundamentals

111

Figures 5.7 and 5.8. Alkyl ammonium ion IPMCs have larger displacement and slower response speed than the alkali metal ion IPMC. From these experimental results, a qualitative model of the IPMC actuator is proposed based on the ion cluster structure shown in Figure 5.9. It is well known that the fluorinated ion exchange polymer swollen with water has a hydrophilic channel-linked ion cluster structure surrounded with a hydrophobic perfluoro backbone polymer network. Counter ions and water transfer through hydrophilic narrow channel-linked ion cluster. If a hydrophilic small cation, such as Liþ, Naþ, Kþ and so on, transfers through the channel, large mobility and little electro-osmosis take place, which result in higher response and smaller displacement. In the case of hydrophobic large cations, such as TEA and TBA, the reverse occurs. It has been reported that the normalized displacement per charge of the Nafion 117/Pt was proportional to the water transference coefficient of the counter cation [18, 19]. These results will be explained quantitatively in the next section based on the micromechanical model of electric responsive ionic gel.

Figure 5.9 Schematic drawing of the ion cluster structure of the fluorinated ion exchange polymers with (a) alkali metal counter ion and (b) alkyl ammonium counter ion forms (Reprinted with permission from Onishi, K., Sewa, S., Asaka, K. et al. The effects of counter ions on characterization and performance of a solid polymer electrolyte actuator, Electrochimica Acta, 46 (8), 1233–41. Copyright (2001) Elsevier).

The structure of the plated electrode also affects the performance of the IPMC actuator [20]. Images of the cross-section of the Flemion/Au composite plated by different plating cycles are shown in Figure 5.10. By adsorption–reduction cycling, a fractal-like structure of gold with high interfacial area within the membrane was obtained. Figure 5.11 shows the plots of the electric double-layer capacitance of the Flemion/Au composites against the number of plating steps, which were estimated by the cyclic voltammogram. The double-layer

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Biomedical Applications of Electroactive Polymer Actuators

100 µm

100 µm

20 µm

20 µm

100 µm

20 µm

20 µm 8 plating cycles

6 plating cycles

4 plating cycles

2 plating cycles

100 µm

Figure 5.10 Scanning electron micrograph of the cross-section for the Flemion/Au composite of different plating cycles (Reprinted with permission from Onishi, K., Sewa, S., Asaka, K. et al. Morphology of electrodes and bending response of the polymer electrolyte actuator, Electrochimica Acta, 46 (5), 737–43. Copyright (2001) Elsevier).

Capacitance (µF cm–2)

2500 No.1

2000

No.2 1500 1000 500 0

0

1

2 3 4 5 6 7 Number of Plating Steps

8

Figure 5.11 Plots of the double-layer capacitance of the Flemion/Au composite of different plating cycles (Reprinted with permission from Onishi, K., Sewa, S., Asaka, K. et al. Morphology of electrodes and bending response of the polymer electrolyte actuator, Electrochimica Acta, 46 (5), 737–43. Copyright (2001) Elsevier).

IPMC Actuators: Fundamentals

113

capacitance increases as the number of plating steps, since the effective electrode area increases, as shown in the SEM images in Figure 5.10. The bending amplitude is proportional to the accumulated charge as shown in the model. Hence, the displacement is proportional to the double-layer capacitance. By the development described above, an IPMC actuator having an optimized performance for a specific application, such as a quick response (Figure 5.12a) and a very large response (Figure 5.12b), can be fabricated.

a. Na+ form

b. (C4Hg)4Na+ form

Figure 5.12 Images of electric response of the Flemion/Au composite actuator in water: (a) Naþ form (b); (C4H9)4Nþ form (b) in water (Reprinted with permission from Onishi, K., Sewa, S., Asaka, K. et al. The effects of counter ions on characterization and performance of a solid polymer electrolyte actuator, Electrochimica Acta, 46 (8), 1233–41. Copyright (2001) Elsevier).

5.5 Model In order to explain the actuation behaviour and develop its performance, many workers have modelled the IPMC actuator. Some have developed a black box model, in which the IPMC actuator is considered as a black box, and the response function for determining the relationship between input and output [21–24]. The black box model is useful for applications. However, we cannot understand the mechanism of the electromechanical response of the IPMC. In this section, a physics-based IPMC model based on the electro-responsive gel theory is introduced. We proposed an IPMC model, in which the bending response is attributed to the electro-osmotic flow in the ion gel film, in 2000 [19], taking into account only the water flow in the membrane. In the same year, de Gennes et al. [25] gave a more comprehensive theory using the phenomenological equation for the electric current ( je) and the water flux ( js) based on the irreversible thermodynamics: je ¼ e r js ¼ rp

lrp lr

ð5:5Þ ð5:6Þ

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Biomedical Applications of Electroactive Polymer Actuators

where e is the conductance,  is Darcy’s permeability and l is Onsager’s coupling constant. The actuation and sensor properties were qualitatively discussed by de Gennes et al., but they have not done any quantitative analysis based on their equations. Yamaue et al. [26] developed the electrostress diffusion coupling (ESDC) model as more formal theory for the deformation dynamics of ion gels under electric field. The theory is a straightforward extension of the stress diffusion coupling model, which was proposed by Doi in 1990 [27]. In their formulation, by discussing the coupling of the network stress with the solvent permeation of the gel network, adding a kinetic equation for the ion flux and electric potential, a complete set of equations for the gel deformation was derived. In their model, the Onsagar coefficients in Equations (5.5) and (5.6) are represented by microscopic parameters: e ¼

l¼

¼

cp q2p X ci q2i þ ; p i i

qp ð1 p ð1

p Þ þ

ð5:7Þ

X ci qi wi ; i i

p Þ 2 X ci w2i þ ; cp  p i i

ð5:8Þ

ð5:9Þ

where suffixes p, s and i denote polymer, solvent and ions, respectively, and c, q, z and w represent the concentration (number of molecules per unit volume), charges, friction constant related to solvent and the specific volume, respectively. The general model was applied to the ion cluster structure of fluorinated ion exchange membrane shown in Figure 5.9. The friction constant of free ions is given by the Stokes– Einstein law:  i ¼ 6pai ;

ð5:10Þ

where  is the viscosity of solvent and ai is the ion radius. To estimate zp, it was assumed that the gel consists of microchannels of characteristic length zb, which correspond to the diameter of the microhydrophilic channel shown in Figure 5.9. Then, the friction constant of polymer gel is given by: p ¼

6p : 2b cp

ð5:11Þ

Using Equations (5.10) and (5.11), the Onsager coefficients are written:   cp q2p b 3 cp  b þ e ¼ 6pb ai cp qp 2b l¼ 6p

(

ð1

p Þ

4p 3



ai b

2 )

ð5:12Þ

ð5:13Þ

IPMC Actuators: Fundamentals

¼

ð1

8 p Þ 2 2b < 1þ : 6p

4p 3ð1

p Þ

!2

cp 3b



ai b

9 5 = : ;

115

ð5:14Þ

Equations 5.12 and 5.13 show that the conductivity decreases and the electro-osmosis increases as the ion size increases in relation to the channel size. This is qualitatively in agreement with the displacement experiments shown in Figures 5.7 and 5.8. The initial curvature is given by the equation: 1 4 l ¼ Q RðtÞ h2 e

ð5:15Þ

R–1(ai/ξb,t = 0)/R–1(ai/ξb = 0.1,t = 0)

which means that initial curvature is driven by the pressure gradient in the ion gel due to the electro-osmotic flow. Hence, the initial curvature is a function of ionic charge (Q) and the water transference coefficient (l/e), which represents how many water molecules transfer per one counter cation transfer. Figure 5.13 shows the normalized value of initial curvature of the IPMC actuator (Nafion 117/Au) of various ionic forms, together with the theoretical curves. Experimental and theoretical values are in good agreement each other.

10 : experiment λ = λi : theory z=1

8

λ = λi z=2

6

λ = λp + λi z=1

5 Li+

λ = λp + λi

2 H+

0

0

0.1

K+

Ca++

Na+

0.3

0.2

z=2

0.4

0.5

ai/ξb

Figure 5.13 Calculated initial curvature of various counter ions to that of proton (Hþ): R(Hþ)/ R(0) (Reprinted with permission from Yamaue, T., Mukai, H., Asaka, K. and Doi, M., Electrostress Diffusion Coupling Model for Polyelectrolyte Gels, Macromolecules, 38 (4), 1349–56. Copyright (2005) American Chemical Society).

One more important conclusion derived from ESDC theory concerns the back motion of the IPMC actuator. As a result of the initial charging of the electric double-layer, a large pressure gradient is created near the electrodes, which results in the initial curvature as given by Equation (5.15). Accordingly, the water starts to flow from the high pressure region to the low pressure region. As a result, the IPMC

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Biomedical Applications of Electroactive Polymer Actuators

starts to bend back. According to Yamaue et al., the characteristic time of the back motion is given by:  relax ¼

h2 ¼ D0 

2

l e

h2 

4 Kþ G 3

ð5:16Þ



D’ is the collective diffusion coefficient of the fluorinated ion exchange polymer, which is given by the effective solvent permeability coefficient and the mechanical constant of the polymer network. The effective solvent permeability coefficient depends on the ratio of the

2

log10 (τrelax/τrelax, H+)

: Experiment : Theory

TEA+

1.5 Z=1 Z=2

1 Ca++

0.5 Na+

0

H+

0

Li+

K+

0.1

0.3

0.2

0.4

0.5

ai/ξb

Figure 5.14 Calculated ratio of the relaxation time for various counter ions to that of proton (Hþ):  relax / relax,Hþ (Reprinted with permission from Yamaue, T., Mukai, H., Asaka, K. and Doi, M., Electrostress Diffusion Coupling Model for Polyelectrolyte Gels, Macromolecules, 38 (4), 1349–56. Copyright (2005) American Chemical Society).

ion radius to the diameter of the microchannel. Hence, the characteristic time of the back motion depends on the counter cation, as shown in Figure 5.14. Experimental and theoretical results are good agreement with each other. If large cations such as TEA and TBA are used as counter cations, the characteristic time is infinitely long. Hence, the back motion of the IPMC of alkyl ammonium form cannot be observed. It is considered that the static response of the IPMC actuator can be attributed to the electrostatic and osmosis effect due to the electric double-layer charging. A theoretical model in consideration with these effects has also been developed [28].

5.6 Recent Developments Recently, applications of room temperature ionic liquids to electrochemical devices are very attractive in both scientific and technological fields. In the area of the EAP

IPMC Actuators: Fundamentals

117

actuators, it is very promising to use ionic liquids as a nonvolatile and high conductivity electrolytes for the ion conductive material of the conductive polymer based actuator [29]. Leo and co-workers [30] have developed a method of using ionic liquids as solvents for the IPMC actuators and sensors in order to use them in air. They successfully fabricated the IPMC actuator using the ionic liquid working in air for a long time. Fukushima et al. [31] have developed a dry actuator that can be fabricated simply through layer-by-layer casting with bucky gel, a gelatinous mixture composed of ionic liquid and a single-walled carbon nanotube. A configuration of the actuator, which has a bimorph structure with a polymer-supported internal ionic liquid layer sandwiched by bucky gel electrode layers, is shown in Figure 5.15.

Figure 5.15 Schematic drawing of a configuration of the bucky gel actuator (Reprinted with permission from Fukushima, T., Asaka, K., Kosaka, A. and Aida, T. Fully Plastic Actuator through Layer-by-Layer Casting with Ionic-Liquid-Based Bucky Gel, Ang. Chem. Int. Ed., 44 (16), 2410–3. Copyright (2005) Wiley-VCH Verlag GmbH).

The actuator film was fabricated by layer-by-layer casting of electrode-layer (singlewalled carbon nanotubes (SWNTs) and ionic liquids (ILs)) and electrolyte-layer (ILs) components in a gelatinous mixture of poly(vinylidene fluoride-co-hexafluoropropylene) (PVdF(HFP)) as a polymer support and a solvent. The actuator can be activated by low voltage (<3 V) and is long-lived upon operation in air. The bucky gel actuator can be used for various applications as well as for IPMC actuators.

5.7 Conclusion In this chapter, the fundamental aspects of the ionic polymer–metal composite (IPMC) actuators have been described. The IPMC actuators have many unique characteristics such as softness, large bending response, low voltage drive, easy forming into any shape, and so on, which are suitable for biomedical applications. In the next chapter, the biomedical applications of IPMC actuators are described.

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References 1. Hamlen, R. P., Kent, C. E. and Shafer, S. N. (1965) Electrolytically activated contractile polymer, Nature, 206, 1149–50. 2. DeRossi, D., Kajiwara, K., Osada, Y. and Yamauchi A. (Eds), Polymer Gels – Fundamentals and Biomedical Applications, Plenum Press, New York, 1991. 3. Oguro, K., Kawami, Y. and Takenaka, H. (1992) Bending of an ion-conducting polymer filmelectrode composite by an electric stimulus at low voltage. J. Micromachine Soc., 5, 27–30. 4. Bar-Cohen, Y. (Ed.) Electroactive Polymer (EAP) Actuators as Artificial Muscles, SPIE Press, 2nd edn, Washington, 2004. 5. Shahinpoor, M. (1992) Conceptual design, kinematics and dynamics of swimming robotic structures using ionic polymeric gel muscles, Smart Materials and Structures, 1, 91–4. 6. Tadokoro, S., Konyo, M. and Oguro, K. (2004) Modeling IPMC for design of actuation mechanisms, in Electroactive Polymer (EAP) Actuators as Artificial Muscles (Y. Bar-Cohen Ed.), SPIE Press, 2nd edn, Washington, 385–427. 7. Nemat-Nasser, S. and Thomas, C. W. (2004) Ionomeric polymer-matal composites, in Electroactive Polymer (EAP) Actuators as Artificial Muscles, (Y. Bar-Cohen Ed.), SPIE Press, 2nd edn, Washington, 171–230. 8. Shahinpoor, M., Kim, K. J. and Mojarrad, M., Artificial Muscles – Applications of Advanced Polymeric Nanocomposites, CRC Press, New York and London 2007. 9. Kim, K. J. and Tadokoro, S. (eds) Electroactive Polymers for Robotics Applications, Springer, London 2007. 10. Yeo, R. S. and Yeager, H. L. (1985) Structural and transport properties of perfluorinated ionexchange membranes, Modern Aspects of Electrochem., 16, 437–505. 11. Abe, Y., Mochizuki, A., Kawashima, T., et al. (1998) Effect on bending behavior of counter cation species in perfluorinated sulfonate membrane-platinum composite. Polym. Adv. Tech., 9, 520–6. 12. Asaka, K., Fujiwara, N., Oguro, K., et al. (2001) State of water and ionic conductivity of solid polymer electrolyte membranes in relation to polymer actuators. J. Electroanal. Chem., 505, 24–32. 13. Asaka, K., Fujiwara, N., Oguro, K., et al. (2002) State of water and transport properties of solid polymer electrolyte membranes in relation to polymer actuastors, Proceedings of the SPIE Conference on Electroactive Polymer Actuators and Devices, San Diego, CA, 4695, 191–8. 14. Fujiwara, N., Asaka, K., Nishimura, Y., et al. (1999) Preparation of gold-solid electrolyte composites as electric stimuli responsive materials, Chem. Mat., 12, 1750–4. 15. Onishi, K., Sewa, S., Asaka, K., et al. (2001) The effects of counter ions on characterization and performance of a solid polymer electrolyte actuator, Electrochim. Acta, 46, 1233–41. 16. Asaka, K., Mori, N., Hayashi, K., et al. (2004) Modelling of the electromechanical response of ionic polymer metal composite (IPMC), Proceedings of the SPIE Conference on Electroactive Polymer Actuators and Devices, San Diego, CA, 5385. 17. Asaka, K. and Fujiwara, N. (2003) Electric deformation response of anion-exchange membrane/ gold composite, Electrochimica Acta, 48, 3465–71. 18. Asaka, K., Oguro, K., Nishimura, Y., et al. (1995) Bending of polyelectrolyte membraneplatinum composites by electric stimuli, Part I. Polym. J, 27, 436–40. 19. Asaka, K. and Oguro, K. (2000) Bending of Polyelectrolyte Membrane-platinum composites by electric stimuli. Part II. Response kinetics, J. Electroanal. Chem, 480, 186–98. 20. Onishi, K., Sewa, S., Asaka, K., et al. (2001) Morphrogy of Electrodes and Bending Response of the Polymer Electrolyte Actuator, Electrochim. Acta, 46, 737–43. 21. Kanno, R., Tadokoro, S., Takamori, T., et al. (1995) Modeling of ICPF (ionic conducting polymer gel film) actuator – modelling of electrical characteristics, Proceedings of the IEEE International Conference on Industrial Electronics Control and Instrumentation, 913–8.

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22. Kanno, R., Tadokoro, S., Takamori, T., et al. (1996) Linear approximate dynamic model of ICPF (ionic conductive polymer gel film) actuator, Proceedings of the IEEE International Conference on Robotics and Automation, 219–225. 23. Newbury, K. M. and Leo, D. J. (2003) Linear constitutive model of ionic polymer bender transducers, Proceedings of the SPIE Conference on Electroactive Polymer Actuators and Devices, San Diego, CA, 5051, 88–99. 24. Nemat-Nassar, S. (2002) Progress of experimental characterization and micromechanistic modeling of actuation of ionic polymer metal composites, Proceedings of the SPIE Conference on Electroactive Polymer Actuators and Devices, San Diego, CA, 4695, 32–41. 25. de Gennes, P. G., Okumura, K., Shahinpoor, M. and Kim, K. J. (2000) Mechanoelectric effects in ionic gels, Europhys. Lett., 50, 513–8. 26. Yamaue, M., Mukai, H., Asaka, K. and Doi, M. (2005) Electrostress diffusion coupling model for polyelectrolyte gels, Macromolecules, 38, 1349–56. 27. Doi, M. (1990) Viscoelastic effect of polymer diffusion, in Dynamics and patterns in complex fluids (eds A. Onuki and K. Kawasaki) Springer, 100. 28. Yamaue, M., Nakamura, K., Taniguchi, T. and Doi, M., (2005) The structure of the electric double layer on the polyelectrolyte gels-electrode interfaces, Polym. Preprints, Japan, 54, 4672–3. 29. Lu, W., Fadeev, A. G., Qi, B., et al. (2002) Science, 297, 983–7. 30. Bennett, M. D. and Leo, D. J. (2004) Ionic liquids as stable solvents for polymer transducers, Sensors and Actuators A, 115, 79–90. 31. Fukushima, T., Asaka, K., Kosaka, A. and Aida, T. (2005) Fully plastic actuator through layerby-layer casting with ionic-liquid-based bucky gel, Angew. Chem. Int. Ed., 44, 2410–3.

6 Active Microcatheter and Biomedical Soft Devices Based on IPMC Actuators Kinji Asaka and Keisuke Oguro National Institute of Advanced Industrial Science and Technology (AIST), Japan

6.1 Introduction The ionic polymer–metal composite (IPMC) has much potential for applications, as mentioned in Chapter 5. Hence, there have been many examples of application research, including biomedical and human affinity applications. The most important aspect is to design an optimized IPMC structure for special applications. To achieve this, it is important how to design moulding the ionic polymer, patterning the plating electrode and connecting the electric wire to the patterned electrode. Examples of the various application research using the IPMC actuators are summarized in Figure 6.1, focusing on designing the structure of the actuator element for transforming the bending motion to other modes of action, such as elliptic, three degrees of freedom (3DOF), round, linear, surface, multifreedom, and so on [1–14]. Described in this chapter, are the important points for fabricating the IPMC actuator device, such as moulding the ionic polymer, patterning the plating electrode, optimization of performance by choosing ions and polymers and making of devices. Then, some examples of our biomedical application research of the IPMC actuators are described, including microcatheter, sheet-type Braille display, underwater microrobot and biped walking robot.

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Biomedical Applications of Electroactive Polymer Actuators

Figure 6.1

Various applications of IPMC actuators.

6.2 Fabrication of the IPMC Device One of the advantages of the IPMC as an actuator material is that it can be easily formed into various shapes, its properties can be engineered and it can potentially be integrated with microelectromechanical systems (MEMS) sensors as control devices to produce smart systems. The fabrication of an IPMC device is shown in Figure 6.2. Forming the ion exchange polymer can be easily done by casting the dispersing solution and evaporating the solvent or hot moulding the thermoplastic resin and changing the ion exchange polymer by hydrolysis. Our scheme for the preparation of the Nafion polymer by casting (a) and hot moulding (b) methods is shown in Figure 6.3. In laboratory, Nafion

Active Microcatheter and Biomedical Soft Devices Based on IPMC Actuators Dispersing solution of perfluorinated ion exchange polymer

123

Thermoplastic resin of perfluorinated ion exchange polymer

Casting

Hot-moulding Ion exchange polymer film, tube, etc. Electrode plating IPMC element Patterning electrode Integrating with sensor and control unit. IPMC device

Figure 6.2 Scheme of the fabrication of an IPMC actuator device.

Casting solution of Nafion

Casting on a flat plate

Evaporating a solvent at 50 °C

Thermoplastic type of Nafion (X = SO2F )

Hot moulding (Ex. Hot pressing at 175 °C)

Hydrolysis SO2F–>SO3– in KOH (0.15)/DMSO(0.35)/H2O (0.5) at 70 °C

Crystallization at 100–170 °C Immersion in fresh water for more than one night Conditioning in HCl/H2O2 solution at 70 °C Immersing in 12 % HNO3 (a)

(b)

Figure 6.3 Scheme for the preparation of the Nafion polymer by casting (a) and hot moulding (b) methods.

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Biomedical Applications of Electroactive Polymer Actuators

films of various thicknesses can be obtained easily. Preparing the Nafion film connecting with other materials can also be done by the same methods. The standard method of electrode plating for the IPMC is chemical plating of gold or platinum as described in Chapter 5. Recently, several physical technologies such as printing or pressing using metal or metal oxide nanoparticles were reported [15, 16]. In order to make an IPMC device that can move in multi degrees of freedom, the plating electrode needs to be patterned. In the laboratory, the plating electrode can be patterned by cutting the plating electrode by PC-controlled mechanical or optical (laser) cutting machine. Finally, a soft and smart device is fabricated by integrating the IPMC actuator with control and senor units. Recently, organic transistor control and sensor technologies have been developed. Hence, all organic, flexible electromechanical devices can be constructed using the IPMC and an organic transistor. In this chapter, such an example is described as a sheet-type Braille display. It was reported that the IPMC has a function of not only an actuator but also a sensor, like a piezo-material [17]. Yamakita et al. [18] studied the sensor property of Nafion/Au and developed the IPMC sensor/actuator system. The response of the IPMC actuator was controlled by the feedback of the same IPMC sensor signal.

6.3 Applications to the Microcatheter Catheter based diagnosis and therapy have become increasingly popular. Conventional catheters cannot move actively and the operations require certain human skill. If the direction of the tip of microcatheter and guide wires can be controlled outside the body, it would be very useful to use catheter based diagnosis and therapy. Such catheters are called ‘active catheters’. In order to develop active catheters, several micro-actuators were applied to the active microcatheter system. The IPMC actuator is soft and driven by low voltage. And it is easy to miniaturize. Hence, the IPMC actuator was applied to the active microcatheter for aneurysm surgery in the brain. Guo et al. [8, 9] developed the microcatheter system with active guide wire. They proposed a prototype model of microcatheter with active guide wire, which has two bending degrees of freedom. Prototype models were 3Fr, 4Fr, 6Fr (1Fr ¼ 1/3 mm) in diameter and consisted of catheter tube and active guide wire with IPMC actuator (Nafion 117 plated with platinum) on its front end as the servo actuator (Figure 6.4). The bending motions and bending angle of the active microcatheter system were measured by application of electricity in physiological saline solution using laser displacement sensors. They also carried out the modelling of this active microcatheter system, which is reasonable for practical applications. By using simulators, they also carried out simulation experiments in vitro using various blood vessel simulators. The system of blood vessel simulators consists of a blood vessel simulator, a pump for circulating physiological saline, an instrument for measurement and a heater (Figure 6.5). The specifications of blood vessel simulator are 4 mm, 5 mm and 8 mm in internal diameter, turning angle of 45–95° and sectional diameter of 2–8 mm in aneurysms. The catheters used for experiments are 6Fr, 4Fr, 3Fr in outer diameter. When the temperature of

Active Microcatheter and Biomedical Soft Devices Based on IPMC Actuators Lead wire

Micro tube

125

Electrode

φ dg

ICPF

Bond Length

Figure 6.4 Prototype model of a microcatheter system with active guide wire using IPMC (ICPF) actuator (Reprinted with permission from Guo, S., Fukuda, T., Kosuge, K. et al. Micro catheter system with active guide wire-structure, experimental results and characteristic evaluation of active guide wire catheter using ICPF actuator, Proceedings of the IEEE 5th International Symposium on Micro Machine and Human Science, 191–8. Copyright (2004) IEEE).

Instrument for measurement

Infusion pump

Heater

Turning point

Aneurysm Simulator Catheter

Figure 6.5 Schematic drawing of a system of blood vessel simulator for testing a microcatheter with the IPMC active guide wire (Reprinted with permission from Guo, S., Fukuda, T., Kosuge, K. et al. Micro catheter system with active guide wire-structure, experimental results and characteristic evaluation of active guide wire catheter using ICPF actuator, Proceedings of the IEEE 5th International Symposium on Micro Machine and Human Science, 191–8. Copyright (2004) IEEE).

physiological saline ranges from 20 to 36 °C, and the flow rate ranges between 50 and 650 ml/min, the inserting motion into each aneurysm and at divergence’s were confirmed as shown in Figure 6.6. According to Guo et al., these experimental results in vitro indicate that the proposed catheter with active guide wire works properly, and that it can improve the effectiveness of traditional procedure for intracavity operations. If the microcatheter itself is moveable, active microcatheter system is more compact and easier to control than the active guide wire. Oguro et al. [10] developed the tubular IPMC actuator with four electrodes and active microcatheter system without guide wire using the tubular IPMC actuator. Figure 6.7 shows a schematic drawing of the tubular IPMC actuator with four electrodes around the tube, which can drive the tubular resin to bend in multiple directions. The four electrodes were fabricated by laser ablation cutting after the chemical plating of a gold electrode. In order to bend the tubular actuator of 0.8 mm outer diameter for 90°, the perfluorocarboxylic acid polymer (Flemion) was used for the ionic gel instead

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Figure 6.6 Bending motion of the active guide wire into aneurysm in blood vessel simulator (Reprinted with permission from Guo, S., Fukuda, T., Kosuge, K. et al. Micro catheter system with active guide wire-structure, experimental results and characteristic evaluation of active guide wire catheter using ICPF actuator, Proceedings of the IEEE 5th International Symposium on Micro Machine and Human Science, 191–8. Copyright (2004) IEEE).

Figure 6.7 Tubular IPMC actuator for controlling active microcatheter (Reprinted with permission from Oguro, K. et al. Proceedings of the SPIE 6th International Symposium on Smart Structures and Materials, Newport Beach, 3669, 64–71. Copyright (1999) SPIE).

of the perfluorosulfonic acid polymer (Nafion). By applying the sequential plating of the gold electrode and optimizing the counter cation as alkyl ammonium cation, we successfully developed a tubular actuator for the active catheter. The motion of the tube to all directions can be controlled with combined signals applied to four electrodes. For practical usage of the tip of active microcatheter for intravascular neurosurgery as shown in Figure 6.8, the tubular actuator of external diameter of 0.8 mm in the swollen state was attached to the end of catheter. The actuator is 2 cm long and the total length of the catheter is 1.5 m. Four conductive layers were made on the outside of the long

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Figure 6.8 Schematic representation of active catheter system using the tubular IPMC actuator for intravascular neurosurgery (Reprinted with permission from Oguro, K. et al. Proceedings of the SPIE 6th International Symposium on Smart Structures and Materials, Newport Beach, 3669, 64–71. Copyright (1999) SPIE).

catheter with gold paste. The tip of the active catheter bent more than 90° within 10 seconds in physiological saline without gas evolution. Intravascular in vivo tests of the active microcatheter with an animal were also carried out.

6.4 Other Applications 6.4.1

Sheet-Type Braille Display

The IPMC actuator is soft, flexible and a low voltage drive. Hence, it is very safe for humans. If the IPMC film can be successfully patterned and integrated, and each microactuator strip can be controlled separately, a human-affinity tactile information communication system can be developed. Someya et al. [19] developed a large-area, flexible and lightweight sheet-type Braille display, integrating the IPMC actuator with their high quality organic transistors. Images of the Braille display are shown in Figure 6.9. An array of the rectangular IPMC actuator is mechanically processed using a numerically controlled (NC) cutting machine to form an array of 12  12 rectangular actuators whose size is 1  4 mm2. A small semi-sphere, which projects upward from the rubber-like surface of the display, is attached to the tip of each rectangular actuator. The effective display size is 4  4 cm2. Each Braille letter consists of 3  2 dots and 24 letters can be displayed. The total thickness and the weight of the entire device are 1 mm and 5.3 g, respectively.

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Figure 6.9 Images of a Braille display: (a) the flexible Braille sheet display using IPMC actuator array; (b) the device assembly; (c) circuit diagram of the Braille sheet display (Reprinted with permission from Kato, Y., Sekitani, T., Takamiya, M. et al. Sheet-Type Braille Displays by Integrating Organic Field-Effect Transistors and Polymeric Actuators, IEEE Trans on Electron. Devices, 54 (2), 202–9. Copyright (2007) IEEE).

Figure 6.10 A cross-sectional illustration of a single Braille dot. An organic transistor is connected to an IPMC actuator with silver paste patterned by a microdispenser (Reprinted with permission from Kato, Y., Sekitani, T., Takamiya, M. et al. Sheet-Type Braille Displays by Integrating Organic Field-Effect Transistors and Polymeric Actuators, IEEE Trans on Electron. Devices, 54 (2), 202–9. Copyright (2007) IEEE).

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Each IPMC actuator is connected to one organic transistor. The circuit diagram of the Braille sheet display is shown in Figure 6.9c. The vertical and the horizontal lines represent bit and word lines, respectively. Figure 6.10 shows the cross-sectional structure of a single Braille cell composed of one transistor and one actuator. When a voltage is applied to the IPMC actuator, the IPMC actuator bends (Figure 6.10). The semi-sphere placed on the actuator rises with the voltage supply and pushes up a rubber-like surface. An organic transistor active matrix is used to address the pop-up dots. Figure 6.11a shows one of the Braille dots moving upwards and downwards. Four Braille letters displayed by the present device are also shown in Figure 6.11b. Someya et al. tested the readability of the sheet-like Braille display using the IPMC. Four visually impaired individuals participated in the reading tests. When the operator input ‘Na’ and ‘Wa’ in the Japanese Braille format, all four individuals were able to recognize the letters correctly.

Figure 6.11 Display operation. (a) Magnified pictures of one Braille dot moving upwards and downwards. The scale is 1 mm. (b) Pictures of the Braille sheet display showing the characters ‘l’, ‘w’, ‘b’ and ‘f’ in the American Braille style (Reprinted with permission from Kato, Y., Sekitani, T., Takamiya, M. et al. Sheet-Type Braille Displays by Integrating Organic Field-Effect Transistors and Polymeric Actuators, IEEE Trans on Electron. Devices, 54 (2), 202–9. Copyright (2007) IEEE).

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This result shows that the generating force and the displacement of this device are sufficient to facilitate reading by the visually impaired. Further, it demonstrates the feasibility of this design, which integrates organic transistors and the IPMC actuators, for realizing Braille sheet displays. The issues that remain to be addressed are the readability and stability of the device. The IPMC actuators are usually operated in wet conditions, while organic transistors degrade easily in moisture and/or oxygen. The straightforward approach toward suppressing such degradation is to employ more sophisticated encapsulation techniques. Alternately, the use of the ionic-liquid-gel based polymer actuator technology, which can be operated in ambient air, is an option. 6.4.2

Underwater Microrobot

An underwater microrobot is very attractive for developed medical practice, both for diagnosis and for surgery. For instance, with medical technology a common application is to perform a delicate surgical operation supported using micromachines, thus avoiding unnecessary incisions. Microrobots can restrict their work to the affected part or the breakdown spot and do not unnecessarily influence their surroundings. Gou et al. [2–4] developed a fish-like underwater microrobot, using the IPMC actuator. Figure 6.12 shows a schematic representation of the structure of the microrobot, which consists of the body made of wood materials shaped as a fish, a pair of fins driven by the IPMC (ICPF) actuator, respectively, connecting the lead wires for supplying the electrical energy to the IPMC actuator. The IPMC actuator used was Nafion 117 plated with platinum or gold. The two fins were driven independently. The direction of swimming can be controlled by the frequency of both fins. Figure 6.13 shows photographs of the microrobot. Its overall size is 45 mm in length, 10 mm in width and 4 mm in thickness. The swimming characteristic of the underwater microrobot was measured by changing the frequency of input voltage from 0.1 to 5 Hz in water and the amplitude of the input voltage from 0.5 to 10 V. The experimental results indicate that Fin

ICPF actuator

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Figure 6.12 Schematic drawing of the structure of a microrobot using the IPMC (ICPF) actuator (Reprinted with permission from Guo, S., Fukuda, T. and Asaka K., A new type of fish-like underwater microrobot, IEEE/ASME Trans on Mechatronics, 8 (1), 136–41. Copyright (2004) IEEE).

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Figure 6.13 Images of the microrobot using the IPMC actuator (Reprinted with permission from Guo, S., Fukuda, T. and Asaka K., A new type of fish-like underwater microrobot, IEEE/ASME Trans on Mechatronics, 8 (1), 136–41. Copyright (2004) IEEE).

changing the amplitude and the frequency of input voltage can control the swimming speed of the underwater microrobot using the IPMC actuator. Guo et al. [20] developed the micropump using the IPMC actuator as shown in Figure 6.14. A photograph of the prototype model of the micropump, which has a size of 10 mm in diameter and 20 mm in length, is shown in Figure 6.15. A flow of 4.5–37.8 ml/min was successfully obtained from the micropump by changing the frequency of the applied voltage. The micropump using the IPMC actuator, as shown in Figure 6.14, is able to make a microflow and is silent for driving, which is suitable for biomedical uses. Nakabo et al. [14] developed a snake-like swimming robot with a patterned-electrode IPMC. The aim is that a snake-like motion sweeps a smaller area than simple bending swimming. Thus, it is suitable for future swimming robots in thin tubes, such as blood vessels, as shown in Figure 6.16. A photograph of the patterned-electrode IPMC for the snake-like swimming robot is shown in Figure 6.17. The IPMC used in this study was a Nafion 117 membrane chemically plated five times with gold. The plated electrode was separated into seven segments by cutting the plated gold using a small hand chisel. For the swimming experiment using the

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One-way valve

(a) Liquid from the Inlet

(b) Liquid from the Outlet

Figure 6.14 Mechanism of a micropump using the IPMC actuator (Reprinted with permission from Guo, S., Nakamura, T., Fukuda, T. and Oguro, K. Development of the micropump using ICPF actuator, Proceedings of the 1997 IEEE International Conference on Robotics and Automation, 1, 266–71. Copyright (1997) IEEE).

Figure 6.15 Image of a micropump using the IPMC actuator (Reprinted with permission from Guo, S., Nakamura, T., Fukuda, T. and Oguro, K. Development of the micropump using ICPF actuator, Proceedings of the 1997 IEEE International Conference on Robotics and Automation, 1, 266–71. Copyright (1997) IEEE).

snake-like IPMC robot, seven connectors and electric wires touched each segment with floats, which prevent the IPMC robot from sinking. The applied travelling wave voltage propels the snake-like IPMC robot in the water. Figure 6.18 shows a photograph of the propulsions of the IPMC snake-like robot. An optimal condition for propulsion was studied by changing the frequency and phase shift of the applied travelling wave voltage. The maximum speed was obtained at the frequency of 2 Hz. The propelling direction of the IPMC (forward or backward) was successfully

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Thin tube Communicator, Patterned (blood vessel etc.) wave generator artificial muscle

Figure 6.16 Concept of a snake-like swimming robot in a thin tube (Reprinted with permission from Nakabo, Y. et al. Biomimetic soft robots using IPMC in Electroactive Polymers for Robotic Applications (eds Kim, K. J. and Tadokoro, S.), 165–98. Copyright (2007) Springer).

Figure 6.17 Patterned IPMC actuator film (Reprinted with permission from Nakabo, Y. et al. Biomimetic soft robots using IPMC in Electroactive Polymers for Robotic Applications (eds Kim, K. J. and Tadokoro, S.), 165–98. Copyright (2007) Springer).

Figure 6.18 Forward and backward propulsions of the snake-like robot using the patternedelectrode IPMC actuator (Reprinted with permission from Nakabo, Y. et al. Biomimetic soft robots using IPMC in Electroactive Polymers for Robotic Applications (eds Kim, K. J. and Tadokoro, S.), 165–98. Copyright (2007) Springer).

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controlled by changing the direction and speed of propagation of waves along the body by advancing and delaying the phases of the sine waves. The maximum speed was obtained at the phase shift of 60°. The snake-like IPMC robot was also successfully turned right or left by applying a saw-tooth wave voltage. Nakabo et al. presented a propulsion model of the snake-like swimming motion to explain results of the propulsion and turning of the patterned-electrode IPMC robot. 6.4.3

Linear Actuators for a Biped Walking Robot

The IPMC actuator basically has a bending motion. In order to apply the robotics, an actuator that has a linear motion (linear actuator) is often very practical. Yamakita et al. [11–13] developed a linear actuator using the IPMC actuator. The proposed linear actuator is composed of many basic units connected in parallel and series so that enough force and displacement can be obtained (Figure 6.19). The elementary unit consists of four IPMC films (Figure 6.20). One side of the unit is formed from a pair of films that are connected by a flexible material or the same thin film. When an input voltage is applied to electrodes on the surface with the anode outside, each membrane bends outwards, then the actuator is constricted. The actuation force and displacement of each unit are small; however, the elementary units can be connected in parallel and series, so the actuator can realize the desired force and displacement. elementary unit

Figure 6.19 Concept of a linear actuator using the IPMC actuator (Reprinted with permission from Yamakita, M. et al., Robotic Application of IPMC Actuators with Redoping Capability in Electroactive Polymers for Robotic Applications (eds Kim, K. J. and Tadokoro, S.), 199–225. Copyright (2007) Springer).

Figure 6.20 Structure of the elementary unit of the proposed linear actuator (Reprinted with permission from Yamakita, M. et al., Robotic Application of IPMC Actuators with Redoping Capability in Electroactive Polymers for Robotic Applications (eds Kim, K. J. and Tadokoro, S.), 199–225. Copyright (2007) Springer).

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uh mh r

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Figure 6.21 Model of a biped walking robot using the IPMC linear actuator (Reprinted with permission from Yamakita, M. et al., Robotic Application of IPMC Actuators with Redoping Capability in Electroactive Polymers for Robotic Applications (eds Kim, K. J. and Tadokoro, S.), 199–225. Copyright (2007) Springer).

The basic characteristics of the elementary unit were studied. The IPMC actuator used in the experiment was Nafion plated with gold. A counter cation in the Nafion was sodium. Though the response of the actuator varies depending on its condition, it was confirmed that the unit whose total length is 40 mm is constricted by 10 mm with a step input voltage of 2.5 V on average. As shown in the chapter on the fundamentals of this technology, the bending characteristics of the IPMC actuator depend on the counter cation in the ionic polymer. Yamakita et al. developed the application of the IPMC linear actuator to a biped walking robot (Figure 6.21) and optimized the performance of the actuator by selecting the counter cation in the ionic polymer.

6.5 Conclusions In this chapter, methods of fabricating IPMC actuator devices and various biomedical applications of IPMC actuators have been described. As examples shown in this chapter, IPMC actuators have much potential for biomedical and human-affinity applications. Though there still remain some issues that must be solved, IPMC actuators are expected to be used in various practical biomedical applications in the near future.

References 1. Bar-Cohen, Y., Leary, S., Shahinpoor, M., et al. (1999) Electro-active polymer (EAP) actuators for planetary applications. Proceedings of the SPIE conference on Smart Structures and Materials, Electroactive Polymer Actuators and Devices, Newport Beach, 3669, 57–63. 2. Guo, S., Fukuda, T. and Asaka, K. (2002) Fish-like underwater microrobot with 3 DOF. Proceedings of the IEEE International conference on Robotics and Automation, 738–43.

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3. Guo, S., Fukuda, T. and Asaka, K. (2003) A new type of Fish-Like Underwater Microrobot., IEEE/ASME Trans on Mechatronics, 8, 136–41. 4. Zhang, W., Guo, S. and Asaka, K. (2006) A new type of hybrid fish-like microrobot, Int. J. Automation and Computing, 3, 358–65. 5. Tadokoro, S., Fuji, S., Fushimi, S., et al. (1998) Development of a distributed actuation device consisting of soft gel actuator elements. Proceedings of the IEEE International Conference on Robotics and Automation, 2155–60. 6. Konyo, M., Tadokoro, S., Takamori, T. and Oguro, K. (2000) Artificial tactile display using soft gel actuators. Proceedings of the IEEE International Conference on Robotics and Automation, 3416–21. 7. Tadokoro, S., Yamagami, S., Ozawa, M., et al. (1999) Multi-DOF device for soft micromanipulation consisting of soft gel actuator elements. Proceedings of the IEEE International Conference on Robotics and Automation, 2177–82. 8. Guo, S., Fukuda, T., Kosuge, K., et al. (1994) Micro catheter system with active guide wirestructure, experimental results and characteristic evaluation of active guide wire catheter using ICPF actuator, 5th International Symposium on Micro machine and Human Science Proceedings, Nagoya, 191–7. 9. Guo, S., Fukuda, T., Kosuge, K., et al. (1995) Micro active guide wire catheter using ICPF actuator, IEEE International Conference on Intelligent Robotics and Systems (IROS 95), Pittsburgh, 2, 172–7. 10. Oguro, K., Fujiwara, N., Asaka, K., et al. (1999) Polymer electrolyte actuator with gold electrodes. Proceedings of the SPIE 6th Annual International Symposium on Smart Structures and Materials, Newport Beach, 64–71. 11. Yamakita, M., Kamamichi, N., Kaneda, Y., et al. (2004) Development of an Artificial Muscle Linear Actuator Using Ionic Polymer-Metal Composites. Adv. Robotics, 18, 383–99. 12. Kamamichi, N., Yamakita, M., Kozuki, T., et al. (2007) Doping effects on robotic systems with ionic polymer-metal composite actuators, Adv. Robotics, 21, 65–85. 13. Yamakita, M., Kamamichi, N., Luo, Z.-W. and Asaka, K. (2007) Robotic application of IPMC actuators with redoping capability, in Electroactive Polymers for Robotics Applications (eds Kim, K. J. and Tadokoro, S.), Springer, London. 14. Nakabo, Y., Mukai, T. and Asaka, K. (2007) Biomimetic soft robots using IPMC, in Electroactive Polymers for Robotics Applications (eds Kim, K. J. and Tadokoro, S.), Springer, London. 15. Akle, B. J., Bennett, M. D. and Leo, D. J. (2006) High-strain ionomeric-ionic liquid electroactive actuators, Sensors and Actuators A, 126, 173–81. 16. Levitsky, I. A., Kanelos, P. and Euler, W. B. (2004) Electromechanical actuation of composite material from carbon nanotubes and ionomeric polymer, J. Chem. Phys, 121, 1058–165. 17. Shahinpoor, M., Bar-Cohen, Y., Simpson, J. O. and Smith, J. (1999) Ionic polymer-metal composites (IPMC) as biomimetic sensors, actuators & artificial muscles – a review, Field Response Polym., American Chemical Society, 25–50. 18. Yamakita, M., Sera, A., Kamamichi, N., et al. (2006) Integrated design of IPMC actuator/ sensor, Proceedings of the 2006 IEEE International Conference on Robotics and Automation, 1834–9. 19. Kato, Y., Sekitani, T., Takamiya, M., et al. (2007) Sheet-type Braille displays by integrating organic field-effect transistors and polymeric actuators, IEEE Trans. on Electron Devices, 54, 202–9. 20. Guo, S., Fukuda, T., Nakamura, T. and Oguro, K. (1997) Development of the Micro pump using ICPF actuator, Proceedings of the 1997 IEEE International Conference on Robotics and Automation, Albuquerque, NM, 266–1.

7 Implantable Heart-Assist and Compression Devices Employing an Active Network of Electrically-Controllable Ionic Polymer– Metal Nanocomposites Mohsen Shahinpoor Biomedical Engineering Laboratories, Department of Mechanical Engineering, University of Maine, Orono, USA.

7.1 Introduction Congestive heart failure (CHF) is the number one killer of people all over the world. In the United States alone, cardiovascular and weak heart-related diseases cost the health care systems over US$ 300 billion each year. Currently there is no viable system to assist a weak heart with its ventricular compression deficiencies. There have been many attempts to build a safe and operational artificial heart in the past which we refrain from listing. However, in recent years there has been a number of advanced and totally implantable left-ventricular assist systems (LVAS), such as the Thoratec’s HeartMateÒ (Figure 7.1a), the Baxter’s NovacorÒ (Figure 7.1b) or Arrow International, Inc.’s LionHeartÒ (Figures 7.1c and 7.1d). All are rather invasive, end-stage, temporary and only help with the left ventricular compression and simply are far from being able to fight this deadly disease of humanity, namely CHF. Recently, with the pioneering work of Shahinpoor (1–20), the idea of helping a weak heart by compressing it from without in a soft, intact and intelligent manner by implantable polymeric artificial muscles has gained acceptance in the United States and international medical communities.

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Figure 7.1 Totally implantable LVASs: (a) Thoratec’s HeartMateÒ II; (b) Baxter’s NovacorÒ; (c) and (d) Arrow International’s Lion HeartÒ.

The cost of cardiovascular diseases and stroke in the United States is estimated at well over US$300 billion/year (http://www.americanheart.org). In most severe cases the patient has only a few days to live and a donor heart will be required for any chance of survival, even though the patient’s body is likely to reject the donated transplanted hearts at a rate of 50%. It is in this spirit that it is proposed to develop a family of minimally invasive (without opening the chest of the patient) and thorascopically implantable, intelligent multi-fingered heart compression/assist systems equipped with soft and resilient electroactive polymeric artificial muscles. Shahinpoor has developed soft ionic polymeric artificial muscles that thrive in the wet and saline environment of the inside of the human body and, in particular, around the myocardium of the heart. Such heart compression/assist systems will be capable of selectively assisting the ventricles, and in particular the left ventricle, of a weak heart to produce more internal pressure and to pump more blood from one or more sides in synchrony with the natural systolic contraction of the ventricle, as well as providing arrhythmia control of the beating heart. Note that the heart weighs between 7 and 15 ounces (200–425 grams) and is a little larger than the size of a fist. By the end of a long life, a person’s heart may have beat (expanded and contracted) more than 3.5 billion times. In fact, each day, the average heart pumps about 2000 gallons (7571 liters) of blood and beats 100 000 times. The human heart is located between the lungs in the middle of the chest, behind and slightly to the left of the breastbone (sternum). A double-layered membrane called the pericardium

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surrounds the heart like a sac. The outer layer of the pericardium surrounds the roots of the heart’s major blood vessels and is attached by ligaments to the spinal column, diaphragm and other parts of the human body. The inner layer of the pericardium is attached to the heart muscle. A coating of fluid separates the two layers of membrane, letting the heart move as it beats, yet still be attached to the body. The heart has four chambers. The upper chambers are called the left and right atria, and the lower chambers are called the left and right ventricles. A wall of muscle called the septum separates the left and right atria and the left and right ventricles. The left ventricle is the largest and strongest chamber in the heart. The left ventricle’s chamber walls are only about a half-inch thick, but they have enough force to push blood through the aortic valve and into the body. Four types of valve regulate blood flow through the heart: The tricuspid valve regulates blood flow between the right atrium and right ventricle. The pulmonary valve controls blood flow from the right ventricle into the pulmonary arteries, which carry blood to the lungs to pick up oxygen. The mitral valve lets oxygen-rich blood from the lungs pass from the left atrium into the left ventricle. The aortic valve opens the way for oxygen-rich blood to pass from the left ventricle into the aorta, the body’s largest artery, where it is delivered to the rest of the body (Figure 7.2a). Electrical impulses from the heart muscle (the myocardium) cause the heart to contract. This electrical signal begins in the sinoatrial (SA) node, located at the top of the right atrium (Figure 7.2b). The SA node is sometimes called the heart’s ‘natural pacemaker’. An electrical impulse from this natural pacemaker travels through the muscle fibers of the atria and ventricles, causing them to contract. Although the SA node sends electrical impulses at a certain rate, heart rate may still change depending on physical demands, stress or hormonal factors.

(a)

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Figure 7.2 Basic operation of the heart (a) and electrical impulses and activities (b).

7.2 Heart Failure Heart failure is a clinical syndrome in which the heart fails to maintain an adequate output, resulting in diminished blood flow and congestion in the circulation in the lung or other

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parts of the body or both. Manifestations of engorged blood vessels in the lungs are referred to as left heart failure, and engorgement of veins and capillaries in other parts of the body are called right heart failure. Common causes of heart failure are high blood pressure, coronary artery disease and rheumatic heart disease. Clinical features may vary considerably. The main symptom of left heart failure is shortness of breath, which often occurs during mild exercise and which may result in periodic sudden attacks that frequently occur at night while the victim is lying flat. As heart failure progresses, shortness of breath becomes more difficult to relieve.

7.3 Background of IPMNCs Recent findings on ionic polymer conductor nanocomposites (IPCNCs) and ionic polymer–metal nanocomposites (IPMNCs) as biomimetic distributed nanosensors, nanoactuators and artificial muscles and electrically controllable polymeric network structures have been presented recently [1–20]. Basically, distributed nanosensing and nanoactuation means that these materials are active down to nano size level. In other words, if they are cut as small as nanoactuators and sensors in a few nanometer range, they will still sense and actuate under a voltage of a few microvolts. Furthermore, methods of fabricating several electrically and chemically active ionic polymeric gel muscles – such as polyacrylonitrile (PAN), poly(2-acrylamido2-methyl-1-propane sulfonic) acid (PAMPS) and polyacrylic-acid-bis-acrylamide (PAAM) – as well as a new class of electrically active composite muscle – such as Ionic Polymeric Conductor Composites (IPCCs) or Ionic Polymer–Metal Composites (IPMCs) made with perfluorinated sulfonic or carboxylic ionic membranes – have been introduced and investigated [7]. Several apparatuses for modeling and testing of the various IPMNC artificial muscles have been described to show the viability of the application of both chemoactive and electroactive muscles. Furthermore, fabrication methods of PAN fiber muscles in different configurations, such as spring-loaded fiber bundles, biceps, triceps, ribbon type muscles and segmented fiber bundles, to make a variety of biomimetic sensors and actuators have also been reported [1–20]. Theories and numerical simulations associated with ionic polymer gel electrodynamics and chemodynamics have also been discussed, analyzed and modeled for the manufactured material. In this chapter the focus is on perfluorinated sulfonic ionic multi-functional materials, as potentially powerful ionic polymers for biomimetic distributed nanosensing, nanoactuation, nanorobotics, nanotransducers for power conversion and harvesting, as well as artificial muscles for medical and industrial applications. It must be noted that widespread electrochemical processes and devices use poly (perfluorosulfonic acid) ionic polymers. These materials exhibit [1–20] good chemical stability, remarkable mechanical strength, good thermal stability and high electrical conductivity when sufficiently hydrated and made into a nanocomposite with a conductive phase such as metals, conductive polymers or graphite. As described elsewhere [7], a number of physical models have been developed to understand the mechanisms of water and ion transport in ionic polymers and membranes. Morphological features influence transport of ions in ionic polymers. These features

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have been studied by a host of experimental techniques including: small and wide angle X-ray scattering, dielectric relaxation and a number of microscopic and spectroscopic studies. The IPMNCs are basically a two-phase system made up of a polar fluid (water) containing ion cluster network surrounded by a hydrophobic polytetrafluoroethylene (PTFE) medium. The integrity and structural stability of the membrane is provided by the PTFE backbones and the hydrophilic clusters facilitate the transport of ions and water in the ionic polymer. These nanoclusters have been conceptually described as containing an interfacial region of hydrated, sulfonate-terminated perfluoroether side chains surrounding a central region of polar fluids. Counter ions such as sodium (Naþ) or lithium (Liþ) are to be found in the vicinity of the sulfonates. It must be noted that the length of the side chains has a direct bearing on the separation between ionic domains, where the majority of the polar fluids resides, and the nonpolar domains. High-resolution NMR of some perfluoroionomer shows an unusual combination of a nonpolar, Teflon-like backbone, with polar and ionic side branches. Liu and SchmidtRohr [21] have obtained the first high-resolution NMR spectra of solid perfluorinated polymers by combining 28 kHz magic-angle spinning (MAS) with rotation-synchronized 19F pulses. Their NMR studies enable more detailed structural investigations of the nanometer-scale structure and dynamics of PTFE based ionomers. It has also been well established [1–21] that anions are tethered to the polymer backbone and cations (Hþ, Naþ, Liþ) are mobile and solvated by polar or ionic liquids within the nanoclusters of size 3–5 nanometers.

7.4 Three-Dimensional Fabrication of IPMNCs In a previous work, Kim and Shahinpoor [11] have reported a newly developed fabrication method that can scale up or down the IPMNC artificial muscles in a strip size of micro-to-centi-meter thickness, using the liquid form of perfluorinated ionic polymers. By meticulously evaporating the solvent (isopropyl alcohol) out of the solution, recast ionic polymer can be obtained. A number of these samples are shown in Figure 7.3. (a)

(b)

Figure 7.3 (a) An eight-finger synthetic muscle. It has a thickness of approximately 2 mm. (b) A coil-type synthetic muscle. This coil type muscle creates a linear actuation motion.

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7.5 Electrically-Induced Robotic Actuation In perfluorinated sulfonic acid polymers there are relatively few fixed ionic groups. They are located at the end of side chains, so as to position themselves in their preferred orientation to some extent. Therefore, they can create hydrophilic nanochannels, so called cluster networks [Gierke and Hsu [22] and Gierke, Munn and Wilson [23]]. Such configurations are drastically different in other polymers, such as styrene/divinylbenzene families that limit, primarily by cross-linking, the ability of the ionic polymers to expand due to their hydrophilic nature. Basically, the cations attract water molecules and thus they separate from the polymer backbone charged pendant groups and gather around them a number of water molecules, thus expanding the network or swelling. Once an electric field is imposed on such a network, the conjugated and hydrated cations rearrange to accommodate the local electric field and thus the network deforms, where in the simplest of cases, such as in thin membrane sheets, spectacular bending is observed (Figures 7.4 and 7.5) under small electric fields such as tens of volts per millimeter.

Figure 7.4 A four-fingered IPMNC compression system in open configuration.

Figure 7.5 Typical deformation of strips (10  80  0.34 mm) of ionic polymers under a step voltage of 4 V.

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2.0 ERI-S1 Muscle (5Pt-1Pt/PVP) Lo = 1 inch input = sine

1.8 1.6

0.1 Hz

1.4

δ/Lo

1.2

0.2 Hz

1.0 0.5 Hz

0.8 0.6 0.4 0.2 0.0 0

1000

2000

3000

4000 E (V/m)

5000

6000

7000

8000

2.0 ERI-S1 Muscle (5Pt-1Pt/PVP) Lo = 1.5 inch input = sine

1.8 1.6

δ/Lo

1.4 1.2

0.1 Hz

1.0

0.2 Hz

0.8 0.5 Hz

0.6 0.4 0.2 0.0 0

1000

2000

3000

4000 E (V/m)

5000

6000

7000

8000

Figure 7.6 Displacement characteristics of an IPMNC, ERI-S1 (d: arc length, Lo: effective cantilever beam length). Lo ¼ 1.0 inch (top) and Lo ¼ 1.5 inch (bottom).

Typical experimental deflection curves are depicted in Figures 7.6. Typical frequency dependent dynamic deformation characteristics of IPMNCs are depicted in Figure 7.7. Once an electric field is imposed on an IPMNC cantilever, in the cantilever polymeric network the hydrated cations migrate to accommodate the local electric field. This creates a pressure gradient across the thickness of the beam and thus the beam undergoes bending deformation (Figures 7.6) under small electric fields such as tens of volts per millimeter.

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Biomedical Applications of Electroactive Polymer Actuators

Tip Deflection in mm Tip Force in gramforce

40

30

Tip Deflection points Tip Force points

20

10

0

0

2

4

6

8

10

Step voltage (v)

Figure 7.7 Variation of tip blocking force and the associated deflection if allowed to move versus the applied step voltage for a 10  50  0.3 mm IPMNC Pt-Pd sample in a cantilever configuration.

Figure 7.7 depicts typical force and deflection characteristics of cantilever samples of IPMNC artificial muscles.

7.6 Distributed Nanosensing and Transduction Shahinpoor [19] has presented a review on sensing and transduction properties of ionic polymer conductor composites. Shahinpoor in 1995 [24] and 1996 [25], and recently [26], reported that IPMNCs by themselves and not in a hydrogen pressure electrochemical cells as reported by Sadeghipour, Salomom and Naogi [27] can generate electrical power like an electromechanical battery if flexed, bent or squeezed. Shahinpoor [24–25] reported the discovery of a new effect in ionic polymeric gels, namely the ionic flexogelectric effect, in which flexing, compression or loading of IPMNC strips in air created an output voltage like a dynamic sensor or a transducer converting mechanical energy to electrical energy. Keshavarzi, Shahinpoor, Kim and Lantz [28] applied the transduction capability of IPMNC to the measurement of blood pressure, pulse rate and rhythm measurement using thin sheets of IPMNCs. Motivated by the idea of measuring pressure in the human spine, Ferrara et al. [29] applied pressure across the thickness of an IPMNC strip while measuring the output voltage. Typically, flexing of such material in a cantilever form sets them into a damped vibration mode that can generate a similar damped signal in the form of electrical power (voltage or current) as shown in Figure 7.8. IPMNC sheets can also generate power under normal pressure. Thin sheets of IPMNC were stacked and subjected to normal pressure and normal impacts

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IPMC Composite Sensor

145

Electrodes

Flip

0.008 0.006

E (volts)

0.004 0.002 0.000 –0.002 –0.004 –0.006 0.0

0.5

1.0

1.5

2.0

2.5

3.0

3.5

Time (s)

Figure 7.8 A typical voltage response of an IPMNC strip (10  40  0.2 mm) under oscillatory mechanical excitations.

and were observed to generate large output voltage. Endo-ionic motion within IPMNC thin sheet batteries produced an induced voltage across the thickness of these sheets when a normal or shear load was applied. A material testing system (MTS) was used to apply consistent pure compressive loads of 200 N and 350 N across the surface of an IPMNC sheet (2  2 cm). The output pressure response for the 200 N load (73 psi) was 80 mV in amplitude and for the 350 N (127 psi) it was 108 mV. This type of power generation may be useful in the heels of boots and shoes or places where there are a lot of foot or car traffic. The output voltage of the thin sheet IPMNC batteries [26] under 200 N normal load is depicted in Figure 7.9. The output voltage is generally about 2 mV/cm length of the IPMNC sheet. IPMNCs also enjoy fairly consistent operation under oscillatory activation, as depicted in Figure 7.10. This characteristics is necessary for cyclic heart compression to the tune of billions of times.

146

Biomedical Applications of Electroactive Polymer Actuators IPMC response under compression (200N load) Sensor Output Load

200

250 200

150

150

100

100

50

50

0

0

–50

0

10

20

30

40

50

60

70

Load (N)

Sensor Output (MV)

250

–50 80

Time (s)

Figure 7.9 Output voltage due to normal 90 ° impact of a 200 N load on a 20  20  0.2 mm IPMNC sample.

8 1 V, 1/2 Hz, Step

7 6 PVP Treated IPMC-Nafion (Li+)

Fg (g)

5 4 3 2

Conventional IPMC-Nafioin (Li+)

1 0 0

1×109

2×109

3×109 4×109 5×109 No. of Cycles

6×109

7×109

8×109

Figure 7.10 Long cycles oscillation of IPMNCs versus blocking force (sample size 5  20  0.2 mm). The environmental chamber was maintained at T ¼ 25 °C and RH ¼ 50–55 %.

7.7 Modeling and Simulation As recently as 2000, Nobel Laureate Pierre de Gennes, Okumura, Shahinpoor and Kim [30] presented the first phenomenological theory for sensing and actuation in ionic polymer–metal composites. Asaka and Oguro [31] discussed the bending of

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polyelectrolyte membrane–platinum composites by electric stimuli and presented a theory on actuation mechanisms in IPMNC by considering the electro-osmotic drag term in transport equations. Nemat-Nasser and Li [32] discussed a modeling on the electromechanical response of ionic polymer–metal composites based on electrostatic attraction/repulsion forces in IPMNCs. Later, Nemat-Nasser [33] proposed a revised version of their earlier paper and stressed the role of hydrated cation transport and mobility within the clusters and polymeric networks in IPMNCs. Nemat-Nasser and Wu [34] have proposed a discussion on the role of the backbone ionic polymer and, in particular, sulfonic versus carboxylic ionic polymers, as well as the effect of different cations such as potassium (Kþ), sodium (Naþ), lithium (Liþ), caesium (Csþ) and some organometalic cations on the actuation and sensing performance of IPMNCs. Tadokoro [35], Tadokoro, Yamagami, Takamori and Oguro [36] and Tadokoro, Takamori and Oguro [37] have presented an actuator model of IPMNC for robotic applications on the basis of physico-chemical phenomena. A recent comprehensive review by Shahinpoor and Kim [12] on modeling and simulation of ionic polymeric artificial muscles discusses the various modeling approaches for understanding the mechanisms of sensing and actuation of ionic polymers and the notion of ion mobility. The underlying principle of IPMNCs actuation and sensing capabilities, which can be described by the standard Onsager formulation using linear irreversible thermodynamics, can now be summarized. When static conditions are imposed, a simple description of mechanoelectric effect is possible based upon two forms of transport: ion transport (with a current density, J , normal to the material) and solvent transport (with a flux, Q ; it can be  assumed that this term is water flux). The conjugate forces include the electric field, E , and  p the pressure gradient, r  . The resulting equation has the concise form of: J ðx; y; z; tÞ ¼  E ðx; y; z; tÞ  L12 r pðx; y; z; tÞ  

ð7:1Þ

Qðx; y; ztÞ ¼ L21 E ðx; y; z; tÞ  K r pðx; y; z; tÞ  

ð7:2Þ





where  and K are the material electric conductance and the Darcy permeability, respectively. A cross coefficient is usually L ¼ L12 ¼ L21. The simplicity of the above equations provides a compact view of the underlying principles of both actuation, transduction and sensing of the IPMNCs. When measuring the direct effect (actuation mode) we work (ideally) with electrodes which are impermeable to ion species flux, and thus Q ¼ 0. This gives: 

L Eðx; y; z; tÞ pðx; y; z; tÞ ¼ r  K 

ð7:3Þ

r  , proportional to r  pðx; y; z; tÞ will, in turn, induce a curvature,   pðx; y; z; tÞ. The relationships between the curvature,   , and pressure gradient, r pðx; y; z; tÞ, are fully  derived and described in de Gennes, Okumura, Shahinpoor and Kim [30]. It should be mentioned that (1/rc) ¼ M(E)/YI, where M(E) is the local induced bending moment and is a function of the imposed electric field E, Y is the Young’s modulus (elastic stiffness) of the strip, which is a function of the hydration H of the IPMNC, and I is the

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Biomedical Applications of Electroactive Polymer Actuators

moment of inertia of the strip. Note that locally M(E) is related to the pressure gradient such that in a simplified scalar format: rpðx; y; z; tÞ ¼ ð2P=t Þ ¼ ðM=IÞ ¼ Y=rc ¼ Y  : 

ð7:4Þ

From Equation (7.4) it is clear that the vectorial form of curvature,   E , is related to the imposed electric field E by:  ¼ ðL=KYÞ E  E

ð7:5Þ

Based on this simplified model the tip bending deflection, max, of an IPMNC strip of length lg should be almost linearly related to the imposed electric field due to the fact that:  ffi ½2  max =ðl2g þ  2max Þ ffi 2  max =l2g ffi ðL=KYÞ E  E

ð7:6Þ

The experimental deformation characteristics depicted in Figures 7.5 and 7.6 are clearly consistent with the above predictions obtained by the above linear irreversible thermodynamics formulation, which is also consistent with Equations (7.5) and (7.6) in the steady state conditions and has been used to estimate the value of the Onsager coefficient, L, to be of the order of 10–8 m2/V-s. Here, a low frequency electric field has been used in order to minimize the effect of loose water back diffusion under a step voltage or a DC electric field. Other parameters have been experimentally measured to be K 10–18 m2/CP,  1 A/mV or S/m. A more detailed set of data pertaining to Onsager coefficient, L, is depicted in Figure 7.11. It must be noted

4 Sample #1 Sample #2 Sample #3

L × 108 [(m/s)/(V/m)]

3

2

1

0 0

5000

10000

15000

20000

25000

E (V/m)

Figure 7.11

Experimental determination of Onsager coefficient, L, using three different samples.

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Figure 7.12 Some movement capabilities of IPMNCs suitable for heart compression.

that IPMNCs are quite capable of undergoing complex movements (Figure 7.12) to accommodate the complex pumping action (systolic–diastolic) of the heart.

7.8 Application of IPMNCs to Heart Compression and Assist in General This chapter discusses the broad category of heart compression and assist and arrhythmia control devices, and more particularly the potential applications of ionic polymer–metal nanocomposite (IPMNCs) biomimetic sensors, actuators and artificial muscles integrated mechatronically as a heart compression device which can be implanted external to the patient’s heart, and partly sutured to the heart, without interfering or contacting with the internal blood circulation. Thus, the proposed IPMNC device can thereby avoid thrombosis and similar complications, which are common to current artificial heart, or heart-assist devices, which may arise when the blood flow makes repeated contacts with nonbiological or nonself surfaces. In compressing a heart ventricle the device must be soft and electronically robust in order not to damage the ventricle. This means that the device should contain control means, such as bradycardic (pacing) and tachyarrhythmic (cardioverting/ defibrillating), to facilitate device operation in synchronism with the left ventricular contraction and capable of transcutaneous recharging of the implanted batteries. The general idea is presented in Figure 7.13. Note that the device is implanted essentially in

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Biomedical Applications of Electroactive Polymer Actuators

42

5 3

30 46

12

10

44

Figure 7.13

General configuration for the proposed heart compression device.

the ribcage of the patient but is supported on a flexible stem that extends to the abdomen. It is also possible to place the supporting structure of the heart compression device on the diaphragm muscle. These details will be worked out during the clinical testing and operation of such devices. The slender flexible stem allows the systolic and diastolic cycles of the heart to continue and yet allows the body of the heat to make swinging motions to one side or the other without any obstacle. In Figure 7.13, 42 is the patient body, 44 is the abdomen area, 46 is the rib cage, 5 is the heart, 3 is the polymeric compression finger made with IPMNCs, 30 is the base of the compression device, 10 is a slender column carrying the electronic wires to the muscle and acting as a flexible support column as well and 12 is the power/battery housing placed in the abdomen. A more detailed drawing of the compression device itself is depicted in Figure 7.14. In Figure 7.14, again 3 denotes the compression fingers made with IPMNCs, 5 is the heart itself, 4 depicts an encapsulated enclosure filled with water to create a soft cushion for the compression fingers, 4d’s are IPMNC based sensors cilia to continuously monitor the compression forces applied to the heart and 3e and 3f are the associated wiring and electronics. Note that, assisting or soft compression of the left ventricle of a weak heart will produce more internal pressure to pump more blood in synchrony with the natural systolic contraction of the ventricle. Additionally, the proposed system will also provide

Implantable Heart-Assist and Compression Devices

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4a

5 4b

4

4d 3 3

4b 3c

3c 4a

4

4c 4c

3e 3d

3f

3f

Figure 7.14

3e

Heart compression device equipped with IPMNC fingers.

arrhythmia control of the beating heart and will be powered by IPMNC artificial muscles for cyclically actuating the resilient compression soft fingers, thereby cyclically pumping blood from the ventricles and allowing the ventricles to refill. These devices will be completely implantable in the body of a patient external of the heart, thereby avoiding thrombogenesis and other complications that may arise from contact between the blood flow and artificial, nonbiological surfaces. The soft fingers comprise suitably mounted platinum and gold electrodes for heart monitoring purposes by means of IPMNC capability to determine the ventricular stroke volume and/or pressure. Specifically, the proposed IPMNC based device will provide entirely electricallycontrollable and micro-processor-controlled multi-fingered resilient sphinctering heart compression devices that can be implanted inside the rib cage of a patient with weak heart and will gently squeeze the weak heart to enhance blood circulation and assist the weak heart. Other configurations are depicted in Figures 7.15 and 7.16.

(a)

(b)

Figure 7.15 Four-fingered heart compression device equipped with thick IPMNCs: (a) before compression; (b) after compression.

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Figure 7.16 The upright configuration of the heart compression device.

The compression devices shown in Figure 7.15 and 7.16 were designed and fabricated from thick (2 mm) ionic polymer–metal composites and were subsequently 24-karat gold-plated. In compressing a heart ventricle the device must be soft and electronically robust in order not to damage the ventricle. This means that the device should contain control means such as bradycardic (pacing) and tachyarrhythmic (cardioverting/defibrillating) to facilitate device operation in synchronism with the left ventricular contraction, and should be capable of transcutaneous recharging of the implanted batteries. Note that the device is implanted essentially in the rib cage of the patient but is supported on a slender flexible stem that extends to the abdomen. The stem allows the systolic and diastolic cycles of the heart to continue and yet allows the body of the heart to make swinging motions to one side or the other without unnecessary restriction. It is also possible to place the supporting structure of the heart compression device on the diaphragm muscle. These details can be worked out during the clinical testing and operation of such devices. Again 3 denotes the compression fingers made with IPMNCs, 5 is the heart itself, 4 depicts an encapsulated enclosure filled with water to create a soft cushion for the compression fingers, 4d’s are IPMNC based sensor cilia to continuously monitor the compression forces applied to the heart and 3e and 3f are the associated wiring and electronics. As designed, this device produces assisting or soft compression of the left ventricle of a weak heart to produce more internal pressure and to pump more blood from one or more sides in synchrony with the natural systolic contraction of the ventricle. Additionally, the system can also provide arrhythmia control of the beating heart. The soft fingers incorporate suitably located electrodes for monitoring the ventricular stroke volume and pressure. A simpler design configuration uses a compression band to assist the heart in its systolic and diastolic cycles of compression– decompression as shown in Figure 7.17. Also, the compression band can be designed such that it can encircle the heart as shown in Figure 7.18.

Implantable Heart-Assist and Compression Devices

Figure 7.17

153

A heart with an IPMNC compression band. (See Color Plate 1).

Figure 7.18 An IPMNC compression band in open and closed configurations.

Presented here are some preliminary data concerning a mini heart compression device equipped with IPMNCs. First measured is the force generated by each strip at 5 V, then measured is the pressure generated when squeezing a small balloon or an animal’s heart (Figure 7.19). Figures 7.20 and 7.21 depict the variation of pressure generated (in mm Hg) with the voltage applied. As discussed before, thick IPMNC polymeric muscles had to be manufactured to generate the required force for heart compression. These were measured experimentally on an animal heart such as a rat’s heart. These newly developed threedimensional IPMNCs have been fully discussed in [37–41] and are briefly described below.

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Biomedical Applications of Electroactive Polymer Actuators

Figure 7.19 Mini heart compression device equipped with IPMNC muscles.

10

9.6

9 Weight (g)

8

7.6

7.4

Surface (dm2)

7

Compression force developed (g)

6 5 4

4

3.6

3.51

3 2

2.7

2.34 1.53

1.17

1 0.28

0.18

1.02 0.42

1.2 0.36

0.55

0.24

0

Thick platinum

Thick gold

Figure 7.20

Very thick platinum

2× thick gold

3× thick gold

Pressure generation versus electrode thickness.

Thin gold

Implantable Heart-Assist and Compression Devices

155

1.4

mmHg developed in the Langedorf

1.2

1 Pressure (mmHg) baseline 74 Pressure (mmHg) baseline 42 0.8

Pressure (mmHg) baseline 22 Pressure (mmHg) baseline 14 T 50 fatigue (min)

0.6

0.4

0.2

0 Thick platinum

Thick gold

Figure 7.21

very thick platinum

2× thick gold

3× thick gold

thin gold

Pressure generation versus electrode thickness.

7.9 Manufacturing Thick IPMNC Fingers The preparation of thick IPMNC fingers follows the procedure outlined in reference [38]. However, the reader is referred to references [14–18] for additional information on IPMNCs. Thus, thick IPMNC strips were manufactured based on the procedure reported in [37] and subsequently equipped with platinum electrodes and gold plating on both sides of the strip with a particle penetration depth of 20 mm. Photographs of the IPMNC samples with the platinum electrodes covered with the gold electrodes are shown in Figure 7.22. It must be mentioned that other types of organ compression, and in particular aortic peristaltic compression to enhance blood circulation and assist a weak heart, are also possible with IPMNC polymeric muscles, as depicted below in Figure 7.23. Furthermore, endoscopic surgical operations with IPMNCs bundled up and insertable through endoscopic incisions is also possible, as depicted in Figure 7.24.

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Biomedical Applications of Electroactive Polymer Actuators (a)

(b)

(c)

Figure 7.22 Manufacturing sequence of the heart compression device: (a) four IPMNC fingers cut to scale; (b) the fingers assembled between two gold ring electrodes; (c) the fingers placed between the ring electrodes and closed.

24

3

Figure 7.23 IPMNCs configured to perform peristaltic compression on aortic blood flow and enhance blood circulation and assist a weak heart.

Implantable Heart-Assist and Compression Devices

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4

3c

4

Figure 7.24 Endoscopically surgically insertable configuration for IPMNC polymeric muscles for heart and organ compression.

7.10 Conclusions Design and development of a novel mammalian heart compression/assist device in the form of a multi-fingered robotic hand equipped with an assembly of thick ionic polymer—metal composite (IPMNC) fingers encased inside a water bag were described. The designed and laboratory-tested multi-fingered heart compression device (MFHCD) was shown to be entirely endoscopically implantable. It was further shown that these devices can be directly or transcutaneously energized by inductive magnetoresonant generator. Therefore, based on the background technologies and their successes and failures in treating weak hearts, it is highly desirable to develop a soft heart compression device for patients with CHF problems. In this connection, ionic polymer–metal composites as soft biomimetic sensors, actuators and artificial muscles present a tremendous opportunity.

References 1. Shahinpoor, M. and Osada, Y. (1995) Heart tissue Replacement with Ionic Polymeric Gels, Proceedings of the 1995 ASME Winter Annual Meeting, San Francisco, CA, 314–8. 2. Shahinpoor, M. (2002) Electrically-Controllable Multi-Fingered Resilient Heart Compression Devices, US Patent Office, US Patent 6,464,655, Issued 15 October 2002. 3. Tozzi, P., Shahinpoor, M., Hayoz, D. and L. von Segesser (2004) Electroactive Polymers to Assist Failing Heart: The Future Is Now, Proceedings of the Second World Congress On Biomimetics and Artificial Muscle (Biomimetics and Nano-Bio 2004), Albuquerque, NM, 5–8 December. 4. Shahinpoor, M., Electrically-Controllable Multi-Fingered Resilient Heart Compression Devices (CIP), (2007) US Patent Office, Patent No.7,198,594, CIP to US Patent 6,464,655, Issued 3 April 2007. 5. Shahinpoor, M. (2004) Artificial Muscles, in Encyclopedia of Biomaterials and Biomedical Engineering (eds Wnek, G. and Bowlin, G.), Marcel Dekker Publishers, NY.

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6. Shahinpoor, M., Kim, K. J. and Mojarrad, M. (2004) Ionic Polymeric Artificial Muscles, ERI/ AMRI Press, Albuquerque, NM. 7. Shahinpoor, M., Kim, K. J. and Mojarrad, M. (2007) Artificial Muscles: Applications of Advanced Polymeric Nano Composites (1st edn), Taylor & Francis Publishers, London. 8. Kim, K. J. and Shahinpoor, M. (2007) Special Issue: Biomimetics, Artificial Muscles, and NanoBio 2004, J. Intelligent Material Systems and Structures, 7 (18), 101. 9. Shahinpoor, M. and Hans-Jo¨rg Schneider (eds), (2008) Intelligent Materials (1st edn), Royal Society of Chemistry (RSC) Publishers, Cambridge. 10. Shahinpoor, M., and Kim, K. J. (2001) Ionic Polymer-Metal Composites – I. Fundamentals, Smart Materials and Structures Int. J., 10, 819–33. 11. Kim, K. J. and Shahinpoor, M. (2003) Ionic Polymer-Metal Composites – II. Manufacturing Techniques, Smart Materials and Structures (SMS), Smart Materials and Structures Int. J., 12 (1), 65–79. 12. Shahinpoor, M. and Kim, K. J. (2004) Ionic Polymer-Metal Composites – III. Modeling and Simulation As Biomimetic Sensors, Actuators, Transducers and Artificial Muscles, Smart Materials and Structures Int. J., 13 (4), 1362–88. 13. Shahinpoor, M. and Kim, K. J. (2005) Ionic Polymer-Metal Composites – IV. Industrial and Medical Applications, Smart Materials and Structures Int. J., 14 (1), 197–214. 14. Choi, K., Kim, K. J., Kim, D., et al. (2006) Performance Characteristics of Electro-Chemically Driven Polyacrylonitrile Fiber Bundle Actuators, J. Intelligent Material Systems and Structures, 17 (7), 563–76, July. 15. Shahinpoor, M. (2005) Soft Plastic Robots and Artificial Muscles, Int. J. Adv. Robotic Systems, 2 (2), 161–74. 16. Shahinpoor, M. (2005) Smart Ionic Polymer Conductor Composite Materials as Multifunctional Distributed Nanosensors, Nanoactuators and Artificial Muscles, Am. Soc. Mech. Eng., Mat. Div. (Publication), MD, 485–9. 17. Shahinpoor, M. (2004) Artificial Muscles, in Encyclopedia of Biomaterials and Biomedical Engineering (eds Wnek, G. and Bowlin, G.), Marcel Dekker Publishers, 43–52, NY. 18. Shahinpoor, M. (2004) Electroactive Ion Containing Polymers, in Hand Book of Smart Systems, Institute of Physics (IOP) Publication, London. 19. Shahinpoor, M. and Guran, A. (2003) Ionic Polymer-Conductor Composites (IPCC) as Biomimetric Sensors, Actuators and Artificial Muscles, Selected Topics, in Structures and Mechatronics Systems (ed Belyaev, A. and Guran, A.), World Scientific Publishers, London, 417–436. 20. Shahinpoor, M. (2003) Ionic Polymer-Conductor Composites As Biomimetic Sensors, Robotic Actuators and Artificial Muscles-A Review, Electrochimica Acta, 48 (14–16), 2343–53. 21. Liu, S. F. and Schmidt-Rohr, K. (2001) High-Resolution Solid-State 13C NMR of Fluoropolymers, Macromolecules, 34, 8416–8. 22. Gierke, T. D., and Hsu, W. Y. (1982) The cluster-network model of ion clustering in perfluorosulfonated membranes, in Perfluorinated Ionomer Membranes (eds Eisenberg, A. and Yeager, H. L.), ACS, Washington, DC, 283–307. 23. Gierke, T. D., Munn, G. E. and Wilson, F. C. (1982) Morphology of perfluorosulfonated membrane products—wide-angle and small-angle X-ray studies, ACS Symp. Ser., 180, 195–216. 24. Shahinpoor, M. (1995) A New Effect in Ionic Polymeric Gels: The Ionic ‘Flexogelectric’ Effect, Proceedings of the SPIE 1995 North American Conference on Smart Structures and Materials, 28 February–2 March, San Diego, CA, 2441, 42–53. 25. Shahinpoor, M. (1996) The Ionic Flexogelectric Effect, Proceedings of the 1996 Third International Conference on Intelligent Materials, ICIM’96, and Third European Conference on Smart Structures and Materials, Lyon, France, June 1996, 1006–11.

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26. Shahinpoor, M. (2004) Smart Thin Sheet Batteries Made With Ionic Polymer-Metal Composites (IPMCs), Proceedings of ASME-IMECE 2004 International Mechanical Engineering Congress and RD&D Exposition, 13–19 November, Anaheim, CA. 27. Sadeghipour, K., Salomon, R. and Neogi, S. (1992) Development of A Novel Electrochemically Active Membrane and Smart Material based Vibration Sensor/damper, Smart Materials and Structures J., 1, 172–9. 28. Keshavarzi, A., Shahinpoor, M., Kim, K. J. and Lantz, J. (1999) Blood Pressure, Pulse Rate, and Rhythm Measurement Using Ionic Polymer-Metal Composite Sensors, Elactroactive Polymers, SPIE (publication number 3669-36), 369–76. 29. Ferrara, L., Shahinpoor, M., Kim, K. J., et al. (1999) Use of Ionic Polymer-Metal Composites (IPMCs) As A Pressure Transducer In the Human Spine, in Electroactive Polymers, SPIE (publication, number 3669-45), 394–401. 30. de Gennes, P. G., Okumura, K., Shahinpoor, M., M. and Kim, K. J. (2003) Mechanoelectric Effects in Ionic Gels, Europhysics Letters, 50 (4), 513–8. 31. Asaka, K. and Oguro, K. (2000) Bending of Polyelectrolyte Membrane Platinum Composites by Electric Stimuli, Part II. Response Kinetics, J. Electroanal. Chem., 480, 186–98. 32. Nemat-Nasser, S. and Li, J. Y. (2000) Electromechanical Response of Ionic Polymer-Metal Composites, J. Applied Phys, 87 (7), 3321–31. 33. Nemat-Nasser, S. (2002) Micro-Mechanics of Actuation of Ionic Polymer-Metal Composites (IPMCs), J. Applied Phys, 92 (5), 2899–915. 34. Nemat-Nasser, S. and Wu, Y. (2003) Comparative Experimental Study of Ionic Polymer-Metal Composites with Different Backbone Ionomers And In Various Cation Forms, J. Applied Phys, 93, 5255–67. 35. Tadokoro, S. (2000) An Actuator Model of ICPF for Robotic Applications On the Basis of Physico-Chemical Hypotheses, Proceedings of the IEEE International Conference on Robotics and Automation, 1340–6. 36. Tadokoro, S., Yamagami, S., Takamori, T., and Oguro, K. (2000), Modeling of Nafion-Pt Composite Actuators (ICPF) by Ionic Motion, in Proceedings of the SPIE 7th Smart Structures and Materials Symposium, EAPAD Conference, San Diego, CA, March 2000, 3987, 92–102. 37. Tadokoro, S., Takamori, T. and Oguro, K. (2001) Application of the Nafion-Platinum Composite Actuator, in Proceedings of the SPIE 8th Smart Structures and Materials Symposium, EAPAD Conference, San Diego, CA, March 2001, 4329, 28–42. 38. Kim, K. W. and Shahinpoor, M. (2001) Development of 3-D Polymeric Artificial Muscles, in Proceedings of the SPIE 8th Smart Structures and Materials Symposium, EAPAD Conference, San Diego, CA, March 2001, paper number 4329–30. 39. Adolf, D., Shahinpoor, M., Segalman, D. and Witkowski, W. (1993) Electrically Controlled Polymeric Gel Actuators, US Patent Office, US Patent 5,250,167, Issued 5 October 1993. 40. Shahinpoor, M. and Mojarrad, M. (2000) Soft Actuators and Artificial Muscles, US Patent Office, US Patent 6,109,852, Issued 29 August 2000. 41. Shahinpoor, M. (1995) Spring-Loaded Ionic Polymeric Gel Linear Actuator, US Patent Office, US Patent 5,389,222, Issued 14 February 1995.

8 IPMC Based Tactile Displays for Pressure and Texture Presentation on a Human Finger Masashi Konyo and Satoshi Tadokoro Graduate School of Information Sciences, Tohoku University, Japan

8.1 Introduction Tactile sensation is an important cue for us to find a subtle difference in handling objects. For a robotic telemanipulation surgery, tactile displays, which produce virtual cutaneous sensation on human hands, provide haptic feedback to the operators for finding diseases carefully and manipulating tools dexterously. For virtual reality applications in medicine, such as rehabilitation and psychotherapy, a tactile display can also contribute to human emotional responses because tactile sensation is highly related to both our comfort and wonder. Many researchers have developed tactile displays and haptic interfaces, as reported by Benali-Khoudja [1] and Hayward [2]. However, it is difficult for conventional tactile displays to synthesize rich and complex tactile sensation arbitrarily. The most characteristic feature of tactile sensation is the variety of perceptual contents reflected, from physical factors such as rigidity, elasticity, viscosity, friction and the surface shape of the target material. It is interesting that tactile receptors in human skin cannot sense the physical factors directly. They only detect the inner skin deformation caused by contact with the object. This fact suggests that a tactile display does not have to reproduce the same physical factors of the material to represent virtual touch. In other words, virtual touch needs only to reproduce the internal deformations in the skin. Furthermore, reproduction of only the nervous activities of the tactile receptors can provide the virtual touch regardless of the inner

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deformation. Based on this standpoint, several researchers have proposed tactile display methods to produce virtual touch by generating stimulation of each tactile receptor selectively. Shinoda [3] developed a tactile display using a magnetic oscillator and air pressure to make selective stimulation. Kajimoto [4] used electrocutaneous stimulation to control nervous excitements. Their methods could represent several complex tactile sensations but the problem remains of how to stimulate all kinds of tactile receptors. In addition, active perceptual processes based on touch by movement by hands are quite important for human tactile perception. Hand movement is used consciously or actively to clarify material properties. Gibson [5] reported that active touch is superior to passive touch in quality and quantity. For the conventional tactile displays, however, it was difficult to perform touch movement freely in a 3-D space due to the limits of the size and weight of the device. Electroactive polymer (EAP) materials have many attractive characteristics as a ‘soft’ and ‘light’ actuator for realizing a compact tactile display. The authors have developed a tactile display using ionic polymer–metal composite (IPMC) actuators [6–9]. We successfully developed a wearable tactile display presenting mechanical stimuli on a finger in response to hand movements [9]. In our research, the target of tactile information is quite different from conventional ones. Our display can synthesize several tactile sensations, such as pressure, roughness and friction [9]. It can also produce rich and complex touch feel, even including qualitative information such as a haptic impression or material feel when we stroke a surface of cloth [6, 9]. In this chapter, we introduce the wearable tactile display and the tactile synthesis method using IPMC actuators. Our display can represent texture feeling and pressure sensation by controlling three physical characteristics: roughness, softness and friction.

8.2 IPMC Actuators as a Tactile Stimulator Conventional tactile displays could hardly control delicate tactile sensation, because it was difficult to make fine distributed stimuli on a human skin under the limitations of their actuators, such as magnetic oscillators, piezoelectric actuators, shape memory alloy actuators, pneumatic devices, and so on. EAP materials have many attractive characteristics as a ‘soft’ and ‘light’ actuator for such a stimulation device. The ionic polymer–metal composite (IPMC, which is also known as ICPF in the robotics field) [10, 11] is one of the electroactive polymers that have shown potential for practical applications. The IPMC is an chemically plated electroactive polymer (EAP) material that bends when subjected to a voltage across its thickness (Figure 8.1). A Nafion–gold composite type IPMC [12], which contained the sodium ion, has a relatively good Electrode Pt or Au layer Bending movement PFS membrane

Figure 8.1 Ionic polymer–metal composite (IPMC) actuator.

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performance in both response speed and displacement. The IPMC generates a relatively small force where a cantilever-shaped actuator (2  10  0.18 mm) can generate about 0.6 mN, and therefore its applications need to be scoped accordingly. Some of the applications that were investigated for IPMCs include an active catheter system [13, 14], a distributed actuation device [15–17], and so on. The IPMC has many advantages for a tactile stimulator, including: 1. High spatial resolution: The required spatial resolution for stimulating sensory receptors, especially Meissner’s corpuscle in the fingertip, is less than 2 mm. IPMC films are easy to shape, and their simple operating mechanism allows a stimulator to be miniaturized to make a high-density distributed structure as shown in Figure 8.2a. Conventional actuators can hardly control such a minute force because of their heavy equivalent mass and high mechanical impedance. An IPMC has enough softness that special control methods are not required to use the passive material property.

Human skin Fingertip

Normal stress Shearing stress

Bending movement

IPMC actuator

IPMC actuators (a) Distributed stimulation

(b) Multi-directional stimulation

Figure 8.2 IPMC actuators as tactile stimulators.

2. Wide frequency range: A tactile display can stimulate several tactile receptors selectively by changing frequency ranges because each tactile receptor has different time response characteristics for vibratory stimulation [9]. The required frequency range is from 5 to 200 Hz to stimulate all kinds of tactile receptors. The response speed of an IPMC is fast enough to make a vibratory stimulation on a skin higher than 200 Hz. This means that an IPMC can stimulate all receptors selectively. 3. Stimuli in multiple directions: Each of the tactile receptors has selectivity for the direction of mechanical stimuli. Meissner’s corpuscle detects especially the shearing stress toward the skin surface. Figure 8.2b shows that bending motions of an IPMC, which contacts with a surface of skin in a tilted position, make a stress in both the normal direction and shearing direction. 4. Wearability: To generate the virtual reality of tactile feel, we should move our hand actively and freely, and receive appropriate stimuli in response to the hand movements. An IPMC based wearable display was successfully developed, which was made so small in size and weight that there was no interference with hand movements [8]. 5. Safety: The low driving voltage (less than 5 V) is safe enough to touch with a human finger directly.

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8.3 Wearable Tactile Display A wearable tactile display was developed using IPMC actuators [8]. Figures 8.3a and 8.3b show a close-up view of the stimulator and of it mounted on a fingertip. The structure of the wearable stimulation device is shown in Figure 8.4. The ciliary part is provided with Au-Nafion composite actuators, where each cilium is 3 mm long and 2 mm wide, in 12 rows leaving 1 mm gaps horizontally and 1.5 mm gaps vertically. All cilia are tilted 45° to transmit mechanical stimuli both in the normal and the tangential directions to the surface of the skin efficiently, as shown in Figure 8.2. The power supply line of the IPMC is provided with a flexible wiring board in to minimize restrictions on the hand, so the fingertip can be bent. The use of silicon rubber of 25  25  8 mm applied to the base of the ciliary part has made it possible to lighten the device to approximately 8 g including the flexible wiring board.

(a)

(b)

Figure 8.3 Overviews of the wearable tactile display: (a) close-up of the IPMC stimulator; (b) mounted on a finger.

Figure 8.4 Structure of the tactile stimulator using IPMC actuators.

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An IPMC needs to be kept moistened because its actuators are operated by ionic migration. Even in slightly moist air, however, the device can provide stimuli sufficiently for several minutes. The total display system is shown in Figure 8.5. The stimulation device is attached to the middle fingertip. The system is designed to read positional information of the hand using Polhemus FASTRAK, which can read information according to a magnetic field.

Figure 8.5 Wearable tactile display system in response to virtual contact motion.

8.4 Selective Stimulation Method for Tactile Synthesis In human skin, tactile receptors generate elementary sensations such as touch, pressure, vibratory sensation, pain, temperature sense, and so on. A tactile impression is an integrated sensation of these elementary sensations. To present tactile feel arbitrarily, stimuli applied to these receptors should be controlled selectively and quantitatively. As mentioned previously, tactile receptors cannot sense the physical factors of environments directly. They detect only the skin deformation caused by contacting objects. A tactile illusion can be provided by the reproduction of the activities of tactile receptors, regardless of the inner deformations. The concepts of the selective stimulation method are illustrated in Figure 8.6. There are four types of mechanoreceptors embedded in human fingers: FA I (Meissner’s corpuscle), SA I (Merkel corpuscle), FA II (Pacinian corpuscle) and SA II (Ruffini endings) [18]. It is known that each receptor has temporal response characteristics for mechanical stimulation and causes subjective sensation corresponding to its responsive deformation. For example, SA I detects static deformations of skin and produces static pressure sensation, and FA I detects the velocity of the deformation and produces the sense of fluttering vibration. Tactile impression is an integrated sensation of these elementary sensations. To present tactile feel arbitrarily, stimuli applied to these receptors should be controlled selectively.

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Figure 8.6 Concept of the selective stimulation method for tactile synthesis.

Figure 8.7

Thresholds of tactile receptors for vibratory stimulus.

The first problem is how to stimulate each receptor selectively. We have focused on the frequency response characteristics of the tactile receptors. Figure 8.7 illustrates the human detection threshold against vibratory stimuli, which represents the sensitivity of each receptor to frequency variation [19]. A smaller amplitude threshold means higher sensibility. This figure shows that there are three frequency ranges in which the most sensitive receptor changes. In the lowest frequency range, SA I is most sensitive, relatively. The best becomes FA I in the middle range and FA II is best in the highest range. This suggests that the selective stimulation can be realized using these frequency characteristics, and arbitrary tactile feels can be produced by synthesizing several frequency components.

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For the IPMC tactile display, selective stimulation is realized by changing drive frequencies, using the receptors’ response characteristics. It was confirmed by a subject’s introspection that the contents of sensation vary with the change of drive frequency as follows: 1. Less than 5 Hz: static pressure sensation (SA I). 2. 10–100 Hz: periodic pressing or fluttering sensation, as if the surface of a finger is wiped with some rough material (FA I). 3. More than 100 Hz: simple vibratory sensation (FA II).

8.5 Texture Synthesis Method We focused on the following three sensations to produce a total textural feeling related to the physical properties of materials: softness, roughness and friction. These sensations are fundamental to express the textural feel of cloth-like materials. The three sensations are produced by the following parameters based on the proposed method described later: 1. Softness sensation: the amount of pressure sensation when the finger contacts the surface (Section 8.6). 2. Roughness sensation: changes in the frequency and the amplitude caused by the relationship of the wavelength of the desired surface and hand velocity (Section 8.7). 3. Frictional sensation: changes in the amount of subjective sensation in response to hand acceleration when the finger slides across a surface (Section 8.8). The problem is how to connect the stimulation on each receptor with contact phenomena caused by hand movements and the physical properties of objects. We have proposed stimulation methods connected to the relationship between hand movements and the physical properties of objects [9]. The softness sensation can also represent concavo-convex surface using as a pressure distribution map as shown in Figure 8.8. When we stroke on the high-pressure area, which

Figure 8.8 Display of concavo-convex surface with texture feel using pressure distribution map.

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is represented by gray-scaled colour, we can feel convex surface. If the tactile stimulation for texture feeling is superposed on the pressure map, we can feel the concavo-convex feel with the texture feel. For the roughness sensation, the frequencies of natural stimuli caused by contacting rough surfaces are changed in response to hand movements. Human beings have the possibility to use those changes of frequencies positively. It is known that the slope of the detection threshold of FA I is –1 in the range of less than 40 Hz. The activities of FA I reflects vibratory frequencies proportionally. This means that FA I can perform as a frequency analyser in a certain range. Based on this hypothesis, we proposed a frequency modulation method for displaying the roughness sensation in response to hand velocity, as described in Section 8.7.

8.6 Display Method for Pressure Sensation 8.6.1

Method

It is known that SA I detects static deformations of the skin and generates a static pressure sensation [18]. Therefore, selective stimulation on SA I can generate pressure sensations. As shown in Figure 8.7, the detection thresholds of SA I have flat frequency characteristics in the range of less than 100 Hz. In most of the range in Figure 8.7, FA I is more sensitive than SA I. However, in the range of less than 5 Hz, SA I becomes more sensitive than FA I. This means that very low frequency vibration can generate pressure sensations relatively larger than the sensation of FA I. The authors confirmed that this assumption was true when the amplitude of simulation was small enough not to sense the vibratory sensation.

8.6.2 Evaluation In this experiment [9], the wearable tactile display system shown in Figure 8.5 was used. The subjects put the device on the right middle finger. They could perform a stroking movement in the horizontal direction. The stimulation was simple sinusoidal vibrations at a frequency from 2 to 5 Hz. The stimulations were generated only when the hand velocity was higher than 25 mm/s despite the direction of movement. To measure the pressure sensation, the subjects pushed their left middle finger on a sponge that was set on an electric balance, controlling their finger to the same amount of the pressure sensation of the artificial pressure sensation for three seconds. Then, the amount of the pressure sensation was calculated as the mean of the force for three seconds. Figure 8.9 shows the relationship between the amplitude of vibration and the amount of pressure sensation at each frequency. The amounts of pressure sensation were calculated by a Z-score because the subjects had different sensitivities for the amount of the subjective sensation. The number in the parenthesis shows the mean value of actual forces at the frequency of 5 Hz as a reference. It was confirmed that as the amplitudes increase, the pressure sensations became larger for every frequency component.

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Figure 8.9 Pressure sensation vs driving voltage of low-frequency stimulation for SA I.

Using this method, the softness of materials, which we feel instantaneously when the finger touches a surface, can be expressed by the parameter of amplitude for the frequency components of 5 Hz. If the pressure sensation is larger, the contacting object has more stiffness.

8.7 Display Method for Roughness Sensation 8.7.1

Method

As mentioned in Section 8.5, we suppose that human beings perceive the roughness sensation as the change in frequency detected by FA I in the relationship between their hand movements and the physical properties of the roughness of materials. The roughness of the surface is defined approximately as a sinusoidal surface, which has a given wavelength l as shown in Figure 8.10. When the finger slides on the sinusoidal surface

Figure 8.10 Definition of surface form using wavelength.

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at a given velocity, v, the frequency of stimuli, f, which are generated in a fingertip, is expressed by a wave equation: f ¼

v l

ð8:1Þ

This equation shows that if the hand velocity becomes faster or if the wavelength, l, becomes smaller, the frequency, f, increases. We should consider the response characteristics of FA I, which is known as a tactile receptor related to the roughness sensation. It is known that FA I responds to the velocity of mechanical stimuli [18]. Here, when the finger slides across the surface, as shown in Figure 8.10, a displacement of stimulus, y, at a given time, t, is defined as a sinusoidal function: y ¼ a sinð2pftÞ

ð8:2Þ

where a is the amplitude of stimulation. Thus, the velocity of stimulation is expressed by substituting Equation (8.1) in the following equation: dy v v ¼ 2pa cosð2p tÞ dt l l

ð8:3Þ

This equation presents the information detected by FA I and shows that both the amplitude and the frequency change in response to the velocity, v. Based on this assumption, the roughness sensation can be presented by changing both the frequency and the amplitude of stimulation in accordance with hand velocity. In this manner, the roughness sensation can be defined by the wavelength, l. For practical use of this method, we applied phase adjustments to produce smooth outputs in response to changing frequencies with respect to each sampling time. The frequencies generated by the proposed method depend on the wavelength and hand movements. However, usual hand movement on the surface of several millimeter wavelengths generates the suitable frequency range for FA 1 consequently.

8.7.2

Evaluation

As evaluation indexes of the roughness sensation, nine kinds of close-set lead balls that had different diameters from 0.5 to 10 mm were used. The wearable tactile display system shown in Figure 8.5 was used. The amplitudes of the stimulations were fixed at 6.0 V ( ¼ the maximum input) and each offset was 0.5 V. The offset was needed to avoid an insensitive zone caused by shortage of amplitudes of the actuators. The subjects put the device on the right middle finger. They touched the index with their left hand at the same time. There was no restriction on time to explore. The subjects were six males in their twenties. Figure 8.11 shows the relationship between the defined wavelengths and the mean value of selected indexes with each error bar representing one standard deviation. The results showed that as the defined wavelength became longer, the roughness sensation seemed to increase when the two half groups were considered separately. Especially, as the

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Figure 8.11 Roughness index selected vs wavelength of the stimuli.

wavelengths became shorter, the standard deviations became smaller and the roughness sensations were expressed clearly. From the results, it was confirmed that roughness sensation could be expressed by the parameter of the wavelength in the case of relatively short wavelengths. In addition to the wavelength, it is confirmed that the maximum amplitude of stimulus affects the amount of the subjective sensation of roughness.

8.8 Display Method for Friction Sensation To express a cloth-like textural feeling in response to contact motions, synthesis of both the roughness sensation and softness sensation is not enough. In this section, we introduce the friction sensation. In this study, the definition of friction sensation is not a usual description based on physical contact conditions. We assumed that the friction sensation can be produced as changes in the amount of subjective sensation in response to hand acceleration when the finger slides across the surface. Especially, the friction sensation is used to express the sticking tendency of materials at the beginning of sliding motion. The authors confirmed that stimulation of high-frequency components corresponding to the acceleration of hand movements could produce a natural sliding feeling [9]. It is known that FA II detects the acceleration of stimuli, and it seems that FA II is related to the detection of hand movements such as by a gyro sensor. Figure 8.12 illustrates the relationshipship between hand acceleration and amplitudes of the high-frequency component. The high-frequency component is fixed at 200 Hz, in which FA II become most sensitive. Therefore, the parameters of the friction sensation are the maximum and minimum values of the amplitude shown in Figure 8.12.

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Figure 8.12 Relationship between the amplitude of high-frequency components for the friction sensation and the acceleration of hand movements.

8.9 Synthesis of Total Textural Feeling 8.9.1

Method

In this section, syntheses of total textural feeling related to the physical properties of materials based on the three methods described above were evaluated. The voltage inputs generated by the three methods were combined into a signal by a simple superposition. Four materials were selected as targets of the tactile syntheses. The artificial textural feelings were tuned subjectively by changing the parameters of the roughness, softness and friction sensations. The tunings of textural feelings were extremely easy compared with the author’s conventional study because each parameter was related to the physical properties of the materials. The following were the properties of the four materials and the tuned parameters: 1. Boa: shaggy, thick, uneven and very rough surface (l ¼ 10, a ¼ 5.0, P ¼ 0.0, Fmax ¼ 2.0) 2. Towel: rough surface, thick and soft (l ¼ 2.0, a ¼ 3.0, P ¼ 2.0, Fmax ¼ 1.0) 3. Fake leather: flat surface, thin, hard and high friction (l ¼ 8.0, a ¼ 1.0, P ¼ 4.0, Fmax ¼ 3.0) 4. Fleece: smooth surface, thin, soft and low friction (l ¼ 0.5, a ¼ 1.0, P ¼ 5.0, Fmax ¼ 1.0) 8.9.2

Experiments

Four artificial textures, which were tuned as mentioned above, were evaluated. The four real materials, that is boa, towel, fleece and fake leather, were used to compare the artificial tactile feels. The wearable tactile display system shown in Figure 8.5 was used. The subjects put the device on the right middle or index finger. They could perform stroking motions with their left hand in the horizontal direction. Before the experiments began, the

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subjects had experience with the four textural feelings only once. The subjects compared each artificial texture with the corresponding real material. They were asked to evaluate the similarity of both feelings at five levels (1: Poor, 2: Fair, 3:Good, 4:Very Good and 5:Excellent). There was no restriction on time to explore the textures. The subjects were divided into two groups: three sight-restricted people (two females in their fifties and one female in her forties) and five ordinary persons (five males in their twenties). The sight-restricted people have more sensitive tactile sensation than ordinary persons. It was expected that the sight-restricted people could evaluate more correctly. Figure 8.13 shows the evaluation results for the sight-restricted people and the ordinary persons, respectively. Both the sight-restricted people and the ordinary persons judged a

Figure 8.13 Evaluation of artificial tactile feeling compared with the real materials.

score of more than three, that is ‘Good’, for the almost all artificial textures. These results demonstrated that the proposed methods could synthesize artificial textural feeling corresponding to the real materials. In addition, the sight-restricted people gave higher evaluations than the ordinary persons so that the synthesized textural feelings had the reasonable reality.

8.10 Conclusions In this chapter, we introduce the wearable tactile display and the tactile synthesis method using IPMC actuators. Our display can represent texture feeling and pressure sensation by controlling three physical characteristics: roughness, softness and friction. These parameters of textural feeling can be measured as physical properties. This means that artificial textural feelings could be synthesized automatically, if the tactile sensors could detect such physical parameters. The authors are also developing the tactile transmission system combining the tactile display and tactile sensors as a master–slave system [20].

References 1. Benali-Khoudja, M., Hafez, M., Alexandre, J. M. and Kheddar, A. (2004) Tactile interfaces: a state-of-the-art survey, Proceedings of the 35th International Symposium on Robotics (ISR 2004), Paris France, 23–26 March, 721–726.

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2. Hayward, V. and MacLean, K. E. (2007) Do It Yourself Haptics, Part-I, IEEE Robotics and Automation Magazine, 14 (4), 88–104. 3. Asamura, N., Yokoyama, N. and Shinoda, H. (1998) Selectively Stimulating Skin Receptors for Tactile Display, IEEE Computer Graphics and Applications, 18 (6), 32–7. 4. Kajimoto, H., Inami, M., Kawakami, N. and Tachi, S. (2004) SmartTouch: Electric Skin to Touch the Untouchable, IEEE Computer Graphics & Applications, 23 (1), 36–43. 5. Gibson, J. J. (1962) Observation on Active Touch, Psychological Review, 69 (6), 477–91. 6. Konyo, M., Tadokoro, S., Takamori, T. and Oguro, K. (2000) Artificial tactile feel display using soft gel actuators, Proceedings of the IEEE Intel. Conference on Robotics and Automation, 3416–3421, San Francisco, April 24–28. 7. Konyo, M., Tadokoro, S., Hira, M. and Takamori, T. (2002) Quantitative Evaluation of Artificial Tactile Feel Display Integrated with Visual Information, Proceedings of the IEEE/RSJ International Conference on Intelligent Robotics and Systems, 3060–5, Lausanne, Switzerland, Sep 30–Oct 4. 8. Konyo, M., Akazawa, K., Tadokoro, S. and Takamori, T. (2003) Wearable Haptic Interface Using ICPF Actuators for Tactile Feel Display in Response to Hand Movements, J. Robotics and Mechatronics, 15 (2), 219–26. 9. Konyo, M., Yoshida, A., Tadokoro, S. and Saiwaki, N. (2005) A tactile synthesis method using multiple frequency vibration for representing virtual touch, Proceedings of the IEEE/RSJ International Conference on Intelligent Robotics and Systems, 1121–7, Edmonton, Canada, August 2–6. 10. Oguro, K., Kawami, Y. and Takenaka, H. (1992) Bending of an ion-conducting polymer filmelectrode composite by an electric stimulus at low voltage, J. Micromachine Society, 5, 27–30. 11. Shahinpoor, M. (1992) Conceptual Design, Kinematics and Dynamics of Swimming Robotic Structures using Ionic Polymeric Gel Muscles, Smart Materials and Structures, 1 (1), 91–4. 12. Fujiwara, N., Asaka, K., Nishimura, Y. et al. (2000) Preparation and gold-solid polymer electrolyte composites as electric stimuli-responsive materials, Chem. Mat., 12, 1750–4. 13. Guo, S., Fukuda, T., Kosuge, K., et al. (1995) Micro catheter system with active guide wire, Proceedings of the IEEE International Conference on Robotics and Automation, 79–84, Nagoya, Japan, May 24–26. 14. Onishi, Z., Sewa, S., Asaka, K., et al. (2000) Bending response of polymer electolyete acutator, Proc. SPIE Smart Structures and Materials 2000: Electroactive Polymer Actuators and Devices (EAPAD), Newport Beach, USA, March 6–8, 121–128. 15. Tadokoro, S., Murakami, T., Fuji, S., et al. (1997) An elliptic friction drive element using an ICPF (ionic conducting polymer gel film) actuator, IEEE Control Systems, 17 (3), 60–8. 16. Tadokoro, S., Fuji, S., Fushimi, M., et al. (1998) Development of a distributed actuation device consisting of soft gel actuator elements, Proceedings of the IEEE International Conference on Robotics and Automation, 2155–60, Leuven, Belgium, May 16–20. 17. Tadokoro, S., Fuji, S., Takamori, T. and Oguro, K. (1999) Distributed actuation devices using soft gel actuators, Distributed Manipulation, K. F. Bo¨hringer, H. Choset (Eds), Springer London, 217–235 (2000). (Kluwer Academic Publishers was absorbed by Springer). ˚ . B. and Johansson, R. S. (1984) Properties of cutaneous mechanoreceptors in the 18. Vallbo, A human hand related to touch sensation, Human Neurobiology, 3, 3–14. 19. Maeno, T. (2000) Structure and Function of Finger Pad and Tactile Receptors, J. Robot Society of Japan, 18 (6), 772–775. (In Japanese). 20. Okamoto, S., Konyo, M., Maeno, T. and Tadokoro, S. (2007) Roughness Feeling Telepresence System on the Basis of Real-time Estimation of Surface Wavelengths, Proceedings of the IEEE/ RSJ International Conference on Intelligent Robots and Systems, 2698–2703, San Diego, USA, Oct 27–Nov 2.

9 IPMC Assisted Infusion Micropumps Il-Seok Park1, Sonia Vohnout2, Mark Banister2, Sangki Lee1,3, Sang-Mun Kim1 and Kwang J. Kim1 1 University of Nevada, USA Medipacs LLC, Tucson, USA 3 Volvo Korea, South Korea

2

9.1 Introduction Ionic Polymer–Metal Composite (IPMC) is an attractive Electroactive Polymer (EAP), which is capable of soft actuation, self sensing and energy harvesting. Among the capabilities of the IPMCs, soft actuators and bio-robotic and/or biomimetic applications constitute especially interesting research fields due to the large bending ability, low driving voltage, easy processing and easy miniaturization of IPMCs. Thus, IPMCs are well considered and adapted to robotic actuators, artificial muscles and miniaturized propulsors. Recently, efforts have been conducted in significant biomedical applications, such as artificial muscles, surgical tools and micro medically-used pumps [1]. Other trials to develop the biomedical applications are currently being carried out in various biomedical disciplines: ophthalmology, proctology and urology [2]. In the future, it is expected that applications for IPMCs will broadly spread not only in small-sized biomedical devices but also in large-scale actuators for naval space’s propulsor, as well as in many industrial applications. In this chapter, a prototype infusion micropump is described, along with the relevant modelling of IPMC based micropumps [3, 4].

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Background of IPMCs

IPMCs are smart materials that exhibit electromechanical (actuator) and mechanoelectrical (sensor) applications. Table 9.1 shows performance properties of state-of-the art IPMCs [5]. They bend quickly under a low voltage, as first reported by Oguro and his co-workers [6]. Later, Abe et al. introduced the important role of existent counter ions and their influence during the bending [7]. Asaka and Oguro introduced a theory of the actuation mechanisms [8]; Shahinpoor and Kim demonstrated that the ionic polymer actuator performance depends on the type of cation [9] and further developed a two-step fabrication method [10] in accordance with their findings. In addition, other groups tried to incorporate various metals as electrode materials to articulate physical properties or electrical responses [11–14]. Table 9.1 Performance properties of typical IPMCs (Reproduced with permission from I-S. Park, K. M. Jung, D. Kim, et al. Physical principles of ionic polymer–metal composites as electroactive actuators and sensors, MRS Bulletin, 33, 3, 190–5. Copyright (2008) MRS) Property

Typical

Estimated actuation speed (strain rate) Estimated bending strain Estimated work density Lifetime

3.3 (% sec 1) 0.5–3.3(%) 5.5 (kJ m 3) Up to 1 million cycles (estimated); 250 000 (experimentally reported)

IPMCs consist of both ion exchange polymers acting as base materials and nano-sized metal layers functioning as resistive and capacitive electrodes. Figure 9.1 shows the crosssectional view of typical platinum IPMC that represents the structure of the IPMC with an electrode layer and deposition/diffusion layer on membrane. The actuation and sensing performance of the IPMCs is highly dependent upon the components of the ion exchange polymer (ionic group and cation) and electrode material. Precious metals (or other conductive mediums), especially platinum and gold, have been adopted for electrode metals [15–21]. Also, there is an increasing interest in IPMC paints (Figure 9.2) [22].

Electrode Layer

Pt deposition/diffusion layer on Nafion

COMPO 15.0kV

X400

10µm

WD 15.2mm

Electrode Layer

Figure 9.1 A cross-sectional view of SEM micromorphology of a platinum IPMC that consists of an electrode and diffusion layer in both surface and membrane polymer as substrate.

IPMC Assisted Infusion Micropumps

177

Figure 9.2 A photograph of spraying Nafion using an airbrush (Reproduced with permission from Park, I.-S., Tiwari, R. and Kim, K. J. Sprayed Sensor Using IPMC Paint, Adv. Sci. Tech., 61, 59–64. Copyright (2008) Trans Tech Publications).

9.3 Miniature Disposable Infusion IPMC Micropumps The availability of safe, disposable and robust infusion pumps for intravenous fluid and drug delivery could provide a significant improvement in both private and public healthcare. An infusion pump should infuse medical fluids, bloods and nutrients into a patient without failure. Thus, there has been a demand for accurately controllable pump systems, from small capacity units such as insulin infusion to large volume feeding suppliers. Small volume infusion pumps, especially, are designed to be portable for use not only in a hospital but also for special purpose likes charity and battlefield use [23]. The former small unit infusion pumps are pressured by human blood pressure using a blood pressure cuff. However, this mechanism leads to serious problems, such as sudden change of the flow rate because the infusion flow rate is only dependent on the patient’s blood pressure. Thus, recently, small volume pumps have been usually operated by an automatically controlled motor or a small embedded system. In addition, for the portable infusion pump, other possible pump power sources have been developed, such as osmotic pressure and spring-powered systems. However, with osmotic power needs it is necessary to change a salt solution bag after finishing infusion and with spring power it is necessary to overcome the limitation of the flow rate for use in various conditions. Furthermore, in order to be a certified infusion operation, it is required that the pump can operate if the power cuts out or is even unplugged and can detect the change of flow rate or pressure even if the flow is blocked or kinked, taken backward and finished when an infusion bag or pump is empty. An IPMC has characteristics both as an actuator and sensor. In addition, it is easy to control the frequency (flow interval) and flow rate by changing of the electric signals with low voltage. We propose a new miniaturized disposal IPMC infusion pump with embedded computer controlling, which can control a micro flow with a low power source, batteries.

178

9.3.1

Biomedical Applications of Electroactive Polymer Actuators

Configuration of the IPMC Infusion Pump

The prototype IPMC infusion pump was designed as shown in Figure 9.3. This prototype pump shows the assemblies for evaluating IPMC actuators. This system can be easily scalable. The cover, machined from transparent acrylic, has a series of ten indentations to secure the chambers between two fluid inlets and outlets. Two types of cover were designed, (a) oblong and (b) spherical shape as shown in Figure 9.4. Total volume for the two types was calculated using SolidWorks drawing software: • Spherical: 0.0243 in3 ¼ 0.398205 mL ¼ 398.21 L • Oblong: 0.0415 in3 ¼ 0.680063 mL ¼ 680.063 L

Figure 9.3 An illustration of the IPMC infusion micropump (Reproduced with permission from Vohnout, S., Kim, S.-M., Park, I.-S. and Banister, M., IPMC-assisted miniature disposable infusion pumps with embedded computer control, Proceedings of the SPIE conference 2007. Copyright (2007) SPIE).

(a)

(b)

Figure 9.4 Two types of pump chamber cover; (a) oblong; (b) spherical (Reproduced with permission from Vohnout, S., Kim, S.-M., Park, I. S. and Banister, M., IPMC-assisted miniature disposable infusion pumps with embedded computer control, Proceedings of the SPIE conference 2007. Copyright (2007) SPIE).

IPMC Assisted Infusion Micropumps

179

An ionic liquid treated platinum IPMC (IL-Pt IPMC) was cut using an Engravelab Laser Software and Venus V-12 carbon dioxide laser system (Figure 9.5) by matching the cover and chamber size. The top and general view of the laserfabricated IL-Pt IPMC with oblong type chamber cover are shown in Figure 9.6. To increase the pumping ability and effectiveness, another IPMC cutting design was made, the interdigitated 20-finger IPMC, which is shown in Figure 9.7 together with its 3-D micrographic view.

2.58 10 5

25

10 63 Laser beam line width: 0.2 (a)

(mm)

(b)

Figure 9.5 (a) Laser cutting machine and (b) a CAD design for interdigitated IPMC (Reproduced with permission from Vohnout, S., Kim, S.-M., Park, I.-S. and Banister, M., IPMC-assisted miniature disposable infusion pumps with embedded computer control, Proceedings of the SPIE conference 2007. Copyright (2007) SPIE). (See Color Plate 2).

(a)

(b)

Figure 9.6 (a) The top and (b) general view of laser-designed micropump IPMC with the oblong type chamber cover (Reproduced with permission from Vohnout, S., Kim, S.-M., Park, I.-S. and Banister, M., IPMC-assisted miniature disposable infusion pumps with embedded computer control, Proceedings of the SPIE conference 2007. Copyright (2007) SPIE).

180

Biomedical Applications of Electroactive Polymer Actuators (a)

(b)

IPMC electrode

Figure 9.7 (a) Pictures of the laser cutting designed interdigitated 20-finger IPMC and (b) its digital 3-D micrographic view (Reproduced with permission from Vohnout, S., Kim, S.-M., Park, I.-S. and Banister, M., IPMC-assisted miniature disposable infusion pumps with embedded computer control, Proceedings of the SPIE conference 2007. Copyright (2007) SPIE).

9.3.2

The Control System

A rudimentary controller for the IPMC infusion pump was designed and developed; it employed a Renesas M16C microcomputer (Figure 9.8). In the employed configuration, the controller was capable of actuating ten cells in various patterns. In order to operate the cell’s fingers in two opposing directions, each cell was actuated by a square wave with 50 % duty cycle (12 second cycle time). The finger cells actuated with approximately 5 V limited to a 10 mA signal under this controller. The polarity can be reversed or neutral (þ, 0, –) to allow control of the speed sequence and time of the actuation cycle.

Figure 9.8 The controller for the IPMC micropump (Reproduced with permission from Vohnout, S., Kim, S.-M., Park, I.-S. and Banister, M., IPMC-assisted miniature disposable infusion pumps with embedded computer control, Proceedings of the SPIE conference 2007. Copyright (2007) SPIE).

IPMC Assisted Infusion Micropumps

9.3.3

181

Performance Testing

For successful infusion performance, an infusion pump has to possess the pressure of about less than 55 kPa and 125 kPa in a vein and for epidural or subcutaneous purposes, respectively [24]. Thus, in order to meet the optimum pressure generation, the oblong pump chamber type was adopted. Two IPMC fingers are arrayed into each oblong section of the pump chamber and this configuration was expected to double the force generated. This 20-finger actuator successfully infused the liquid and sequenced with one and two at a time. For achievement of the entire sequence time of the pump within 1–3 s, we are still continuing to sequence the finger actuation program down to fractions of a second. Calculations estimate that the chamber volume can be displaced every three seconds and that we have to subtract the volume of at least one section for each sequence cycle. It would be calculated to 9  0.068 ml ¼ 0.612 ml per cycle time. Therefore, the pump should produce a flow rate of approximately 12.24 ml/min at 1 psi. 9.4

Modelling for IPMC Micropumps

Diaphragms are widely used to create stroke volumes [25, 26]. IPMCs are new, promising materials used for actuating diaphragms in micropumps. In this modelling, systematic design methods on an IPMC actuator-driven valve-less micropump are introduced [3, 4]. IPMC with a Nafion membrane was considered the best material for an actuating diaphragm. In order to estimate deformed shapes of a circle-shaped IPMC diaphragm, the equivalent bimorph beam model [27] for IPMC actuators was used in conjunction with the finite element method (FEM). Using this model, several parametric studies were performed to determine an optimal electrode shape of the IPMC diaphragm and to investigate the pressure effect on the stroke volume. 9.4.1

Equivalent Bimorph Beam Model for IPMC Actuators

To predict the behaviour of IPMC diaphragms, the equivalent bimorph beam model, which was recently introduced by Lee et al. [28], is adopted in this study. Here, the key ideas of the model are summarized. In the equivalent bimorph beam model (Figure 9.9), it is assumed that an IPMC has two virtual layers of the same thickness. Under an imposed electric field across the IPMC, the

Figure 9.9 Typical shape of a bimorph beam (Reproduced with permission from Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications).

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upper layer and the lower layer of an IPMC expand or contract, opposing each other, to produce the IPMCs bending motion. Equation (9.1) shows a relationship between the input voltage (V) and the induced tip displacement (s) of an IPMC with an equivalent bimorph beam: s¼

3d31 VL2 2H 2

ð9:1Þ

where d31 (m/v) is the ‘effective’ electromechanical coupling coefficient for small deflections in which the subscripts 1 and 3 stand for the x-direction and the z-direction, respectively. From Equation (9.1), d31 is expressed as follows: d31 ¼

2sH 2 3VL2

ð9:2Þ

Substituting the experimentally measured tip displacement into Equation (9.2), d31 can be obtained for a given input voltage. In the equivalent bimorph beam model, the Young’s modulus (E) contributing to the bending stiffness of an IPMC is determined from the blocking force Equation (9.3) of a bimorph beam: Fb1 ¼

3WH 2 E d31 E3 8L

ð9:3Þ

where Fb1 is the measured blocking force, E3 is the electric field and d31 is previously calculated from Equation (9.2). To determine the equivalent of Young’s modulus (E), Equation (9.3) is rewritten as: E¼

8LFb1 3WH 2 d31 E3

ð9:4Þ

For all numerical analyses, a commercial finite element analysis program, MSC/ NASTRAN [29], was used in conjunction with the equivalent bimorph beam model. A thermal analogy technique proposed by Taleghani and Campbell [30] was used to implement the electromechanical coupling effect into the finite element model. In the thermal analogy technique, the electromechanical coupling coefficient (d31) is converted into the thermal expansion coefficient 1 as follows: 1 ¼

d31 t

ð9:5Þ

where t is the thickness of a layer across which an electric potential is applied, and then the electric potential (DV) is replaced by the temperature difference (DT). More details and verifications for the thermal analogy technique can be found in Lee et al. [27] and Lim et al. [31]. 9.4.2

IPMC Diaphragm

Deformations of circle-shaped IPMC diaphragms were analysed for the circle-shaped and ring-shaped electrode, respectively. Through parametric studies, an electrode shape was chosen for the optimal diaphragm, which generates maximum stroke volume. In order to show the effectiveness of the circle-shaped diaphragm, its stroke volume was compared to

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183

that of a square-shaped diaphragm maintaining the same actuator area. Both the normal mode analysis and pressure effect on the selected IPMC diaphragm were introduced. 9.4.2.1

Circle-Shaped Electrode vs Ring-Shaped Electrode

Parametric studies on two kinds of electrodes for circle-shaped diaphragms with a radius of 10 mm were conducted with the material properties and thicknesses shown in Table 9.2 [3]. The material properties E and d31 of the IPMC in Liþ form were determined thorough the equivalent bimorph beam model [28]. The elastic modulus of Nafion in Liþ form and Poisson’s ratios were obtained from literature [32, 33]. Figure 9.10 shows the shape of the two electrodes in a one-quarter finite element model of diaphragms. Figure 9.10a is the circle-shaped electrode and 9.10b is the ring-shaped electrode. Table 9.2 Material properties and thicknesses for an IPMC diaphragm (Reproduced with permission from Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC Actuatordriven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications) Elastic Modulus (GPa)

Poisson’s Ratio

*

þ

IPMC in Li form Nafion in Liþ form

1.158 0.05

d31 (m/V)

0.487 0.487

1.750  10 N/A

t (mm) 7

0.2 0.2

* This IPMC is heavily loaded with platinum (~6 vol-% Pt). The Pt loading technique was uniquely designed to enhance the humidity control of IPMC.

Sym.

Sym. Fixed B.C. along the edge

Fixed B.C. along the edge

Sym.

Sym.

Radius of electrode

Radius of electrode

Gray part: IPMC (or electrode) and black part: Nafion (a)

(b)

Figure 9.10 Shapes of electrodes for IPMC diaphragms (¼ finite element model): (a) circleshaped; (b) ring-shaped (Reproduced with permission from Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications).

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Biomedical Applications of Electroactive Polymer Actuators

The total number of elements (Quad4, MSC Software Corp.) [29] used for each model was 400. The symmetry boundary condition was applied to the vertical and horizontal lines, and fixed boundary condition to the outside edge. As shown in Figure 9.10, each IPMC diaphragm consists of the IPMC part and a Nafion part. Therefore, when a voltage is applied on an IPMC part, the vertical interface between IPMC and Nafion can rotate easily to produce large bending deformation, since Nafion has a low elastic modulus. Under an applied 2 V input and fixed boundary conditions along the outside edge, the centre displacements of the diaphragms were calculated with variations of the electrode length in the radial direction. The calculated results are provided in Figure 9.11. For the diaphragm with the circle-shaped electrode, in which the radius of electrode was 8.5 mm, the maximum centre displacement was 0.996 mm. The maximum centre was only 0.686 mm for the diaphragm with the ring-shaped electrode and is more efficient than the ring-shaped electrode in terms of deformation. The parametric studies suggest that there is an optimal radius and radial length to each electrode for the maximum deflections.

1 Center displacement (mm)

Center displacement (mm)

1 0.8 0.6 0.4 0.2 0

0

2

4

6

8

10

0.8 0.6 0.4 0.2 0

0

2

4

6

8

Radius of electrode (mm)

Radial length of electrode (mm)

(a)

(b)

10

Figure 9.11 Centredisplacement of IPMC diaphragms for each electrode case: (a) circle-shaped electrode; (b) ring-shaped electrode (Reproduced with permission from Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications).

For the two optimal electrode cases (radius, 8.5 mm for the circle-shaped electrode; radial length, 5.5 mm for the ring-shaped electrode), stroke volumes were calculated from the diaphragm with the deformed shapes shown in Figure 9.12. Note that the diaphragm with the circle-shaped electrode is bent upward while the diaphragm with the ring-shaped electrode is bent downward for the same electrical input of 2 V. Considering 2 V AC input, calculated stroke volumes (also the definition of stroke volume later in Figure 9.16) for the circle-shaped and the ring-shaped electrode cases were 216 and 104 mL, respectively.

IPMC Assisted Infusion Micropumps

MSC.Patran 2001 r2a 16-Aug04 10:12:28 Fringe: SC1:DIAPHRAGM, A14:Static Subcase: Displacements, Translational-(NON-LAYERED) (ZZ) Deform: SC1:DIAPHRAGM, A14:Static Subcase: Displacements, Translational

185

9.66–004 9.02–004 8.38–004 7.73–004 7.09–004 6.44–004 5.80–004 5.15–004

9.66–004

0.

4.51–004 3.87–004

+

3.22–004 2.58–004 1.93–004 1.29–004 6.44–005

Z

–1.16–010 Y

default_Fringe: Max 9.66–004 @Nd1 Min 0. @Nd 316 default_Deformation: Max 9.66–004 @Nd1

X

(a)

MSC.Patran 2001 r2a 16-Aug04 10:30:37 Fringe: SC1:DIAPHRAGM, A11:Static Subcase: Displacements, Translational-(NON-LAYERED) (ZZ) Deform: SC1:DIAPHRAGM, A11:Static Subcase: Displacements, Translational

0 –4.57–005 –9.14–005 –1.37–004 –1.83–004 –2.29–004 –2.74–004 –3.20–004

0. +

–3.66–004 –4.11–004 –4.57–004

–0

–5.03–004 –5.48–004 –5.94–004 –6.40–004 Z

–6.86–004 Y X

(b)

default_Fringe: Max 0. @Nd316 Min –6.86–004@Nd1 default_Deformation: Max 6.86–004 @Nd1

Figure 9.12 Deformed shapes of IPMC diaphragms: (a) circle-shaped electrode (radius of electrode ¼ 8.5 mm); (b) ring-shaped electrode (radial length of electrode ¼ 5.5 mm) (Reproduced with permission from Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications). (See Color Plate 3).

186

Biomedical Applications of Electroactive Polymer Actuators

9.4.2.2

Circle-Shaped Diaphragm vs Square-Shaped Diaphragm

The shape effect of the diaphragm on the stroke volume was also investigated. A squareshaped diaphragm was modelled and analysed to calculate its centre displacement and stroke volume. The results of the square-shaped diaphragm were compared with those of the circle-shaped diaphragm. The areas of the IPMC and Nafion were maintained the same as for the circle-shaped diaphragm – the optimal case (i.e., the radius of diaphragm: 10 mm and the radius of electrode: 8.5 mm). Material properties and thicknesses are shown in Table 9.2. Figure 9.13 shows the shapes of the two diaphragms (¼ finite element model). For the finite element modelling, 400 and 324 elements (Quad4, MSC.Software Corp.) were used for the circle-shaped diaphragm and the square-shaped diaphragm, respectively. Sym.

Sym.

Sym.

Sym. Gray part: IPMC (or electrode) and black part: Nafion (a)

(b)

Figure 9.13 Shapes of IPMC diaphragms (¼ finite element model): (a) circle-shaped; (b) squareshaped (Reproduced with permission from Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications).

Under a 2 V input, for the square-shaped diaphragm’ the calculated centre displacement and stroke volume were 0.760 mm and 196 mL, respectively. Note that the calculated values for the circle-shaped diaphragm are 0.996 mm and 216 mL, respectively. From the results, it is evident that the use of the circle-shaped diaphragm is advantageous over the square-shaped one in order to generate larger stroke volumes. 9.4.2.3

Normal Mode Analysis

The normal mode analysis was performed for the optimal circle-shaped diaphragm with the circle-shaped electrode (radius of electrode: 8.5 mm). For the calculation, the density of Nafion in Liþ form was 2.078  103 kg/m3 [34] and that of IPMC in Liþ form was assumed to be 2.5  103 kg/m3. Figure 9.14 shows the first and second mode shapes of the diaphragm. The computed first (i.e., fundamental) and the second natural frequencies

IPMC Assisted Infusion Micropumps MSC.Patran 2001 r2a 20-Aug04 14:27:41 Fringe: SC1:DIAPHRAGM, A2:Mode 1: Freq. = 429.69: Eigenvectors, Translational-(NON-LAYERED) (MAG) Deform: SC1:DIAPHRAGM, A2:Mode 1: Freq. = 429.69: Eigenvectors, Translational

187

3.15+002 2.94+002 2.73+002 2.52+002 2.31+002 2.10+002 1.89+002 1.68+002

3.15+002

1.47+002

0.

1.26+002

+

1.05+002 8.41+001 6.31+001 4.21+001 2.10+001

Z

4.20–005 default_Fringe: Max 3.15+002 @Nd1 Min 0. @Nd 316 default_Deformation: Max 3.15+002 @Nd1

Y X

(a)

MSC.Patran 2001 r2a 20-Aug-04 14:28:51 Fringe: SC1:DIAPHRAGM, A2:Mode 2: Freq. = 1659.1: Eigenvectors, Translational-(NON-LAYERED) (MAG) Deform: SC1:DIAPHRAGM, A2:Mode 2: Freq. = 1659.1: Eigenvectors, Translational

3.02+002 2.81+002 2.61+002 2.41+002 2.21+002 2.01+002 1.81+002 1.61+002

02 3.02+0

0.

1.41+002 1.21+002

+

1.01+002 8.04+001 6.03+001 4.02+001 2.01+001

Z

3.81–006 Y X

(b)

default_Fringe: Max 3.02+002 @Nd261 Min 0. @Nd316 default_Deformation: Max 3.02+002 @Nd261

Figure 9.14 Normal mode analysis results for an IPMC diaphragm (radius of electrode ¼ 8.5 mm): (a) first mode; (b) second mode (Reproduced with permission from Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications). (See Color Plate 4).

188

Biomedical Applications of Electroactive Polymer Actuators

are 430 and 1659 Hz, respectively. If we consider the driving frequency range of the IPMC diaphragm as less than 40 Hz, the calculated fundamental frequency is much higher than the driving frequency range. Therefore, the resonance will not affect the stroke volume in that driving frequency range. The results imply that we can linearly control the flow rates of an IPMC-driven micropump within the driving frequency (40 Hz) of interest. 9.4.2.4

Pressure Effect on Stroke Volume

The external pressure effect on the circle-shaped diaphragm having a circle-shaped electrode was investigated. The external pressure could be considered as the chamber pressure of a pump. In order to calculate numerically, the stroke volume under external pressure, uniform pressures were applied to the FE model for the optimal IPMC diaphragm as shown in Figure 13a. Figure 9.15 shows the estimated half stroke volumes of the optimal circle-shaped diaphragm under the upward pressures and 2 V input. In Figure 9.15, ‘Down Stroke’ indicates the half stroke volume when the diaphragm bends downward and ‘Up Stroke’ indicates the half stroke volume when the diaphragm bends upward, as shown in Figure 9.16 for the definition of half stroke volume. According to these results, in the case of the down stroke, the IPMC diaphragm could generate a stroke volume under the upward external pressure up to approximately 2300 Pa, which lies in a range of dynamic pressures of micro air vehicles (MAVs).

Figure 9.15 Half stroke volumes of the circle-shaped IPMC diaphragm (Reproduced with permission from Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-Mass-Flux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications).

IPMC Assisted Infusion Micropumps

189

Aslot v

t2 t0 t1 ∆Vh_us

v = Flow (or jet) speed Aslot = Area of slot ∆Vh_us = Half up-stroke volume ∆Vh_ds = Half down-stroke volume ∆V = ∆Vh_us + ∆Vh_ds = Stroke volume

∆Vh_ds

Figure 9.16 Schematic of an IPMC-driven micropump (Reproduced with permission from Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC Actuator-driven Zero-Net-MassFlux Pump for Flow Control, J. Intelligent Mat. Systems and Structures, 17, 6, 533–41. Sage Publications).

9.5 Conclusions In this chapter, we have described an IPMC-driven infusion micropump for recent biomedical applications. Even though the applications of IPMCs for biomedical fields require more trials and development methods, IPMCs are still attractive materials due to their electromechanical and mechanoelectric properties. A systematic design method of an IPMC-driven micropump was introduced. In order to properly estimate the deformed shapes of IPMC diaphragms, the equivalent bimorph beam model for IPMC actuators was conveniently used, in conjunction with the finite element method.

References 1. Shahinpoor, M. and Kim, K. J. (2005) Ionic polymer–metal composites: IV. Industrial and medical applications, Smart. Mat. Struct., 14, 197–214. 2. Kim, K. J. and Tadokoro, S. (2007) Electroactive Polymers for Robotic Application: Artificial Muscles and Sensors, Springer, London. 3. Lee, S., Kim, K. J. and Park, H. C. (2006) Modeling of an IPMC actuator-driven zero-net-massflux pump for flow control, J. Intelligent Mat. Systems and Structures, 17, 533–9. 4. Lee, S. and Kim, K. J. (2006) Design of IPMC actuator-driven valve-less micropump and its flow rate estimation at low Reynolds numbers, Smart Mat. Struct. 15, 1103–9. 5. Park, I-S., Jung, K. M., Kim, D., et al. (2008) Physical principles of ionic polymer–metal composites as electroactive actuators and sensors, MRS Bulletin, 33, 3, 190–5. 6. Oguro, K., Kawami, Y. and Takenaka, H. (1992) Bending of an ion-conducting polymer filmelectrode composite by an electric stimulus at low voltage, J. Micromachine So., 5, 27–30. 7. Abe, Y., Mochizuki, A., Kawashima, T., et al. (1998) Effect on bending behaviour of counter cation species in perfluorinated sulfonate membrane-platinum composite, Polym. Adv. Techn., 9, 520–6. 8. Asaka, K. and Oguro, K. (2000) Bending of polyelectrolyte membrane platinum composites by electric stimuli Part II. Response kinetics, J. Electroanal. Chem., 480 186–98.

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9. Shahnpoor, M. and Kim, K. J. (2000) Effect of counter-ions on the performance of IPMCs, Proceedings of the SPIE 7th International Symposium on Smart Materials and Structures, Newport Beach, CA, 5–9 March 2000, 3987-18, 110–20. 10. Shahinpoor, M. and Kim, K. J. (2001) Ionic Polymer–Metal composites: I. Fundamentals, Smart Ma. Struct., 10 819–33. 11. Sadeghpour, K., Salomon, R. and Neogi, S. (1992) Development of a novel electrochemically active membrane and smart material based vibration sensor/damper, Smart Mat. Struct., 1, 172–9. 12. Guo, S., Nakamura, T. and Fukuda, T. (1996) Design and characteristic evaluation of micropump using ICPF actuator, Proceedings of the 7th International Symposium on Micro Machine and Human Science, Nagoya, Japan, 2–4 October 1996, 235–40. 13. Shahinpoor, M., Bar-Cohen, Y., Simpson, J. O. and Smith, J. (1998) Ionic polymer–metal composites (IPMCs) as biomimetic sensors, actuators and artificial muscles-a Review, Smart Mat. Struct., 7, R15–30. 14. Dogruer, D., Tiwari, R. and Kim, K. J. (2007) Ionic polymer metal composites as energy harvesters, Proceedings of the 2007 SPIE International Symposium on Smart Materials and Structures, San Diego, CA, 19–22 March 2007, 6524, 1C1–10. 15. Bennett, M. D. and Leo, D. J. (2003) Manufacture and characterization of ionic polymer transducers employing non-precious metal electrodes, Smart Mat. Struct., 12, 424–36. 16. Shahinpoor, M. and Kim, K. J. (2000) The effect of surface-electrode resistance on the performance of ionic polymer–metal composite (IPMC) artificial muscles, Smart Mat. Struct., 9, 543–51. 17. Nemat-Nasser, S. (2002) Micromechanics of actuation of ionic polymer–metal composites, J. Appl. Phys, 92, 2899–915. 18. Kim, S. M. and Kim, K. J. (2008) Palladium buffer-layered high performance ionic polymer– metal composites sensors and actuators, Smart Mat. Struct., 143, 343–351. 19. Lee, D. Y., Park, I.-S., Lee, M.-H., et al. (2007) Ionic polymer–metal composite bending actuator loaded with multi-walled carbon nanotubes, Sensors and Actuators A, 133, 117–27. 20. Bennett, M. D. and Leo, D. J. (2004) Ionic liquids as stable solvents for ionic polymer transducers, Sensors and Actuators A, 115, 79–90. 21. Park, I.-S. and Kim, K. J. (2007) Multi-fields responsive ionic polymer–metal composite, Sensors and Actuators A, 135, 220–8. 22. Park, I.-S., Tiwari, R. and Kim, K. J. (2008) Sprayed Sensor Using IPMC paint, Adv. Sci. Techn., 61, 59–64. 23. Vohnout, S., Kim, S.-M., Park, I.-S., et al. (2007) IPMC assisted miniature disposable infusion pumps with embedded computer control, Proceedings of the 2007 SPIE International Sympoium on Smart Materials and Structures, San Diego, CA, 19–22 March 2007, 6524, 1U1-7. 24. From http://en.wikipedia.org/wiki/Infuion_pump, Infusion pump. 25. Laser, D. J. and Santiago, J. G. (2004) A review of micropumps, J. Micromech. Microeng., 14, R35-64. 26. Woias, P. (2005) Micropumps – past, progress and future prospects, Sensors and Actuators B, 105, 28–38. 27. Lee, S., Park, H. C. and Kim, K. J. (2005) Equivalent modelling for ionic polymer–metal composite actuators based on beam theories, Smart Mat. Struct., 14, 1363–8. 28. Lee, S., Park, H. C., Kim, K. J. and Yoon, K. J. (2004) Equivalent beam and equivalent bimorph beam models for ionic polymer–metal composite actuators, J. Control, Automation, and Syst. Eng., 10, 1012–6. 29. MSC. Software Corp., MSC/NASTRAN user’s manual (2001). 30. Taleghani, B. K. and Campbell, J. F. (1999) Non-linear finite element modelling of THUNDER piezoelectric actuators, NASA/ TM, 209322. 31. Lim, S. M., Lee, S., Park, H. C., et al. (2005) Design and demonstration of a biomimetic wing section using a lightweight piezo-composite actuator (LIPCA), Smart Mat. Struct., 14, 496–503.

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32. Park, H. C., Kim, K. J. Lee, S. and Chah, Y. J. (2004) Electromechanical flapping produced by ionic polymer–metal composites, Proceedings of the 2004 SPIE International Symposium on Smart Materials and Structures, San Diego, CA, March 2004, 5385–63. 33. Nam, J. D., Lee, J. H., Lee, J. H., et al. (2005) Water uptake and migration effects of electroactive IPMC (Ionic Polymer–Metal Composite) actuator, Sensors and Actuators A, 118, 98–106. 34. Nemat-Nasser, S. and Li, J. Y. (2000) Electromechanical Response of Ionic Polymer–Metal Composites, J. App. Phys, 87, 3321–31.

Section III Conjugated Polymers

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

10 Conjugated Polymer Actuators: Fundamentals Geoffrey M. Spinks, Gursel Alici, Scott McGovern, Binbin Xi and Gordon G. Wallace ARC Centre of Excellence for Electromaterials Science and Intelligent Polymer Research Institute, University of Wollongong, Australia

10.1 Introduction Inherently conjugated (or conducting) polymers (ICPs) are one of the main categories of electroactive polymers (EAPs) and fall within the ‘ionic’ category of EAPs. ICPs such as those shown in Scheme 10.1 have been extensively studied for a wide range of applications that use both their inherent conductivity (e.g. sensors [1] or electrostatic discharge protection [2]) and their facile electroactivity: their ability to be electrochemically switched between different states at low voltage with very large changes in properties. The conductivity can change by 10 orders of magnitude; the polymer’s colour changes; the polymer can switch from hydrophilic to hydrophobic; permeability to chemical species changes; the volume changes as does the mechanical properties (e.g. Young’s modulus). These property changes are then useful in a wide variety of devices including electrochromic displays [3], controlled release systems [4], membranes [5] and, of course, actuators. ICPs have been developed for a wide range of actuator applications, as reviewed recently [6, 7]. Compared with other actuator materials, ICPs produce relatively large strains and low/medium stresses and operate at low voltages. Two broad categories of ICP actuators have been developed: linear actuators and benders. The latter are produced when the ICP is

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

196

Biomedical Applications of Electroactive Polymer Actuators R +

+ A– N H

H N

A– S

n

Scheme 10.1

A–

n

m

m

m

Polypyrrole (PPy)

H N

Polythiophene (PTh)

Polyaniline (PANi)

Some simple ICPs (A represents a dopant anion species).

laminated onto a passive, flexible substrate [8]. The actuator system then works like a bimetallic strip so that volume changes in the ICP induce bending. Linear actuators are formed from free-standing films or fibres where actuation in one direction is principally used to perform work. One alternative variant of the linear actuator is those systems that use the thickness direction dimension changes of ICP coatings to affect the mechanical response [9]. One practical example of the latter is a pump system that uses concentric cylindrical layers of ICPs to generate a pumping pressure [10]. Companies dedicated to the development of artificial muscles based on conjugated polymers have also emerged in recent years. MicroMuscle based in Sweden and EAMEX from Japan are both actively pursuing actuators for biomedical and electronics applications. Santa Fe Science and Technology, USA, has produced continuous spun polyaniline fibres and demonstrated their use as linear actuators [11, 12]. Academic laboratories have also developed several demonstration products including a variable camber hydrofoil [13], a robotic fish propulsor fin [14], a gas valve [15], ‘microrobots’ [16] and a micropump [10], some of which are illustrated in Figure 10.1. An electronic Braille screen using ICP actuators is also described in this book [17]. Actuation in ICPs occurs through their reversible redox chemistry. The electrochemical reaction of polypyrrole, one of the most widely studied conjugated polymers, is illustrated Nitrogen

Oxygen cell

Amplifier

3 cm

+ Polymer Actuator

– Air flow

(a)

(b)

(c)

Figure 10.1 Some example devices using conjugated polymer actuators: (a) miniature pump system using concentric polypyrrole tubes; (b) prototype pectoral fin for biorobotic fish (Reproduced from Bioinspir. Biomim, 2, S6-S17, The application of conducting polymers to a birobotic fin propulsor by James Tangorra et al. with permission. Copyright (2007) Institute of Physics Publishing Limited); (c) a ball valve operated by a bender type actuator (Reproduced with permission from Sensors and Actuators A: Physical, 114, 1, Andrews, M.K., Jansen, M.L., Spinks, G. and Wallace, G.G. An integrated electrochemical sensor-actuator system. Copyright (2004) Elsevier; Photograph courtesy Dr Murray Jansen).

Conjugated Polymer Actuators: Fundamentals

197

in Scheme 10.2. The redox switching leads to the addition or removal of charge from the polymer backbone and, thus, the transport of ions into or out of the polymer to balance the charge. These ion movements are principally responsible for the volume changes that produce an actuation response. The extent and speed of actuation in ICPs, therefore, must be influenced by the electrochemical properties of the polymer, as well as the size of the ions that move into and out of the polymer and the speed at which the ions move. These influences will be explored in more detail in subsequent sections. Useful actuators must do work on their surroundings: either moving an attached load and/or generating a force that operates on an attached mechanism. Thus, the actuator material itself is subject to mechanical stress, so the mechanical behaviour of the ICP will also be important in determining the actuation behaviour. The ICP breaking strength limits the maximum stress that can be applied or generated, while the Young’s modulus of the material determines the extent to which the ICP deforms. The mechanical aspects of ICP actuator devices will also be further explored below. Finally, actuators must produce the desired movement accurately and repeatably. Thus, there is a need to develop control systems that ensure that the correct input stimulus is applied to achieve the desired output. Various approaches to modelling the control of ICP actuation are also summarised in this chapter. H N

H N N H

H N N H

N H

X

reduced state

+2e–/–2A–

–2e–/+2A–

oxidised state H N

H N N H

H N N H

–

A

Scheme 10.2

N H

A–

X

The electrochemical reaction of polypyrrole.

10.2 Molecular Mechanisms of Actuation in ICPs The electrochemistry in ICPs occurs by the application of small voltages (typically <1 V) when the ICP is in contact with an electrolyte medium and connected to an auxiliary electrode (which can also be an ICP). As synthesized, via oxidation of the monomer, ICPs are present in the charged conducting form. The charge on the polymer backbone is counter balanced by what is known as the molecular dopant (A ). This dopant can be a simple

198

Biomedical Applications of Electroactive Polymer Actuators

anion such as chloride. That is, for a PPy.Cl film, reduction of the material would result in chloride ion expulsion. In that case and for thin uniform films of the material the reaction described by Equation (10.1) (and Scheme 10.2) would hold true. Assuming this reaction was carried out in chloride-containing media, then re-oxidation would result in chloride ion incorporation. However, consider what happens when A is a large anionic species. Now upon reduction the large dopant anion is immobile and the charge balance is maintained by incorporation of a cation (Mþ) from the electrolyte media in which the reduction occurs. This process is illustrated by Equation (10.2). The preference for an ICP to undergo redox reactions through anion or cation exchange (or a combination of both) is complicated and involves a number of experimental factors, as discussed in more detail below. +

0 +e

A– N H

n

+ mA–

N H

–e m

n

ð10:1Þ

m

+

0

A–

+

N H

mM+

+e

A – M+

ð10:2Þ

N H

–e

m

m

While, fundamentally, the actuation response in ICPs is driven by their electrochemical reactions, a number of other processes can be identified as significantly affecting the actuation performance. The sequence of processes that are known to occur when a potential is applied to the ICP is summarized in Figure 10.2. The important properties of the ICP and of the electrolyte medium are also shown in the figure. Steps 1 and 2 shown in Figure 10.2 more or less occur simultaneously and result in the change in redox state of the polymer. Upon application of the electrochemical potential sufficient to cause a redox reaction in the polymer, electrons are transferred between the Chemical/Mechanical Processes involved in Actuation

1

2

3

4

5

Charge Transfer

Change in polymer oxidation state

Charge balance with ions in electrolyte

Volume Change

Actuation

Ionic conductivity

Polymer conductivity

– Porosity of polymer – Electrolyte/Solvent

Modulus of polymer

Geometry

Material Properties of Polymer and Electrolyte

Figure 10.2 Schematic illustration of some of the molecular level processes that occur when an electrochemical potential is applied to an ICP when in contact with an electrolyte. The material properties involved in each step are also highlighted.

Conjugated Polymer Actuators: Fundamentals

199

polymer and the auxiliary electrode through the external circuit. As electrons flow, the surface of both electrodes becomes charged and ions in the electrolyte solution align themselves at each electrode to balance this charge. The rate and extent that these charge transfer processes can occur is determined by the respective conductivities: ionic conductivity of the electrolyte and electronic conductivity of the polymer. (Any resistance to electron transfer coming from resistance in the external circuit is considered negligible.) The fundamental time constant of these processes is given by the RC time of the system, where R is the total resistance and C the total capacitance. The resistance is inversely proportional to conductivities, so higher conductivities will lead to faster processes. The capacitance will always be large in these systems, since the capacitance is the amount of charged stored (per volt) from electrons and ions transferred to the electrodes. Since the actuation depends on this charge transfer, the capacitance will be large. Step 3 in Figure 10.2 is shown as a separate process, although it represents the continuation of the processes identified in Steps 1 and 2. The fundamental difference with Step 3 is that the electronic charge on the polymer backbone and the associated dopant ion is located within the three-dimensional structure of the polymer and not at the polymer/ electrolyte interface. Ions from the electrolyte must diffuse into the bulk polymer material to balance the charge on the backbone and to enable charge neutrality. The rate and extent that this process occurs is directly related to the diffusion time of ions within the polymer. Hence, the nature of the ionic species, particularly their size, is a major determinant of the speed of Step 3. In addition, the geometry and porosity of the electrode material can also greatly affect the diffusion distances involved and, therefore, the response times. Step 4 shown in Figure 10.2 emphasizes that the mass and charge transport processes cause a volume change within the polymer. The extent of the volume change depends on the amount and size of the ions moved into or out of the polymer. In addition, there are sometimes large quantities of solvent that are also transferred between the polymer and the electrolyte [18]. Finally, it is also recognized that conformational changes to the polymer chains can also occur, which affect the overall volume. The extent that the polymer may swell or contract when an external stress is applied will also be affected by the Young’s Modulus of the material. Changes in modulus during redox cycling can make an additional contribution to the actuation [19]. Finally, the actuation response is represented in Step 5 of Figure 10.2. The way in which the actuation is manifest depends on the geometry of the polymer and the design of the mechanical system. All the above processes (and possibly others) occur simultaneously when an ICP is electrochemically stimulated. The four main processes and how they interrelate are illustrated in Figure 10.3. The complexity of the processes involved means that it is not yet possible to define the optimal ICP composition and geometry to maximize actuation in a given application. In fact, in some cases the structure is modified to improve one process, but the same modification makes another process more difficult. For example, aligning polymer chains to increase electronic conductivity leads to better charge transport but lower actuation strains in the alignment direction [20, 21]. The properties of ICPs vary greatly with the choice of dopant and synthesis conditions. The behaviour of ICP actuators is also very sensitive to the electrolyte used and other environmental conditions. Considerable advances in actuator performance have occurred in recent years, and these advances are reviewed in the following sections with reference to the processes identified in Figures 10.2 and 10.3.

200

Biomedical Applications of Electroactive Polymer Actuators Ionic Conductivity

Electrolytic Salt

Ion Movement

Change in polymer oxidation state

Charge Transfer

Electrolyte concentration

Volume Change

Solvent Solvent movement/ Osmosis

Diffusion Rate Polymer Conductivity

Swelling

Porosity/Morphology

Figure 10.3 The four main processes leading to a volume change and actuation in ICPs, with the main polymer/electrolyte properties illustrated.

10.3 Comparison of Actuation Performance in Various ICPs The largest actuation strains in ICPs have been obtained from polypyrroles (PPy) and these polymers are the most studied. Both polyaniline (PANi) and polythiophene (PTh) ICPs have also been investigated and the solution-processability of these polymers offers some advantages in fabricating actuators. For example, continuous fibres of doped PANi that can be readily bundled to give large force actuators have been prepared [22] (Figure 10.4).

Figure 10.4 Complete polyaniline actuator assembly consisting of eight PANi twisted fibre yarns inserted in a hollow PANi fibre (counter electrode) containing liquid electrolyte. A porous separator prevents short circuiting (Reproduced from Australian Journal of Chemistry, Electrochemical actuator devices based on polyaniline yarns and ionic liquid electrolytes by Lu, W., Norris, I.D. and Mattes, B.R., 58, 263–9. Copyright (2005) CSIRO Publishing).

Conjugated Polymer Actuators: Fundamentals

201

Table 10.1 Actuation performance and various properties of selected ICPs. A web based resource is also available for comparing ICP and other actuator materials [74]

Actuation strain (%) Actuation stress (MPa) Conductivity (S/cm) Reduction potential (V) (vs Ag/AgCl) Elastic modulus (MPa) (wet state) Strain:Charge Ratio (%/C/m3) Data taken from reference:

PPy (high strain)

PPy (typical)

PANi

PTh

35 1 40 NA NA NA [52]

2 5 100 0.5 100–700 8  10 8 [19, 42, 72, 73]

1.5 5 400 þ0.1 NA 6  10 [60]

2 NA NA NA NA NA [74]

9

The actuation performance and other properties typical of PPy, PANi and PTh are summarized in Table 10.1. There is no clear correlation between the properties typical of these polymers and their resultant actuation performance. PPy, for example, shows slightly higher conductivity but similar capacitance to PANi and PTh but also shows vastly higher actuation strains, at least in some formulations. The comparison given in Figure 10.3 highlights the complex nature of actuation occurring in ICPs. The simple picture of redox induced ion exchange between the polymer and the electrolyte inducing volume changes in the polymer is not sufficient to quantitatively describe the actuation. The following sections explore in more detail the mechanisms of actuation in ICPs, starting with their electrochemistry. Specific studies relating to the actuation of polypyrrole are then reviewed, since PPy is the most widely studied ICP actuator.

10.4 Electrochemistry of ICPs Fundamentally, the actuation response in ICPs is driven by their electrochemical reactions. Consequently, it is important to summarize the nature of these reactions and the factors that influence the extent and speed of the molecular processes occurring. In this section, the redox reactions in ICPs are first described followed by a discussion of the factors that determine the extent and rate of reaction that occurs when ICPs are activated. It is shown that oxidation of the polymer produces a positively charged material. The positive charges on the polymer backbone must be balanced by the incorporation of anions, which are also referred as dopant or counterions. During reduction of the polymer, the negative charges encourage anions to leave the polymer backbone. This redox reaction with anion incorporation/ejection is widely revealed using small counterions, such as hexafluorophosphate (PF6 ) [17], perchlorate (ClO4 ) [23], p-phenolsulfonate (PPS ) [24] and tetrafluoroborate (BF4 ) [25]. However, in some cases, where the anions are bulky and immobilized (e.g. dodecyl benzene sulfonate (DBS ) and poly(vinyl phosphate) (PVP )), the charge compensation may be achieved by cation movement from electrolyte. The redox processes predominated by anion incorporation/ ejection and cation incorporation/ejection are shown in Equations 10.1 and 10.2, respectively. In some polymer/electrolyte systems, a mixture of both processes occurs with anion incorporation/cation ejection during oxidation and cation incorporation/anion ejection during reduction [26, 27].

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Biomedical Applications of Electroactive Polymer Actuators

The electrochemical properties of conjugated polymers can be characterized using cyclic voltammetry, the most convenient and reliable electrochemical technique. This method involves the measurement of current at the working electrode as a function of potential during the application of the triangular potential waveform. The current flow is a result of the oxidation/reduction processes and the concomitant ion flow that occurs in ICPs. Cyclic voltammetry provides a rapid determination of electrochemical transitions occurring, the potentials at which these occur, and the rate of these transitions. It is the most effective and versatile electrochemical technique available for the study of redox reactions. Typically, cyclic voltammetry of conjugated polymers is conducted in an electrolyte using three-electrode system with polymer as working electrode together with a counter electrode and a reference electrode. A typical cyclic voltammogram of an ICP is shown in Figure 10.5. As the potential is scanned from negative to positive, there is a net anodic current which increases exponentially with potential corresponding to the electron loss and anion incorporation or cation ejection. At the anodic peak (pa), the redox potential is sufficiently positive where the reduced state polymer is rapidly switched to its oxidized state. Further increases in potential see a decrease in current, which is related to the depletion of oxidizable species near the electrode, so that further oxidation becomes mass-transport controlled. Upon reversal of the scan, the current continues to decay with time until the potential near the redox potential, where the polymer in its oxidized state is switched to its reduced state and a cathodic peak (pc) occurs. The voltammogram is characterized by a peak potential, Epa (peak anodic potential) or Epc (peak cathodic potential), at which the current reaches the maximum value, ipa (peak anodic current) or ipc (peak cathodic current).

0.2 pa

Current (mA)

0.1

0.0

–0.1

–0.2 –1.5

pc

–1

–0.5

0

0.5

1

1.5

2

Potential (V vs Ag/Ag+)

Figure 10.5

A typical cyclic voltammogram of a conjugated polymer.

The sharp peaks in the cyclic voltammogram can be interpreted as electrode potentials of one particular composition of polymer/ions according to the Nernst equation [28]:

Conjugated Polymer Actuators: Fundamentals

RT ared ln nF aox

ox þ n
e ! red ) E ¼ E0

203

ð10:3Þ

where E is the electrode potential, E0 is the standard electrode potential measured at standard ambient conditions, that is at a temperature of 298 K, solutes at a concentration of 1 M and gases at a pressure of 1 Bar. ared and aox are the activities of the reduced species and oxidized species. Activities can be replaced by electrode surface concentrations of the redox couple species, C, to give: E ¼ E0

RT Cred ln nF Cox

ð10:4Þ

This equation indicates that the electrode potential is affected by temperature and the concentrations of the redox species. Large concentration differences can cause the potential to differ significantly from the standard cell potential, and that such large concentration differences also make the potential more temperature dependent. At 25 °C, the expression can also be expressed as: E ¼ E0

0:059 Cred log n Cox

ð10:5Þ

In the case of anion dominated redox processes of conjugated polymers, the reaction can be expressed as: PPyþ A þ e ! PPyþ þ A

) E ¼ E0

0:059 log

CPPyþ
CA ¼ E0 CPPyþ A

0:059 log CA ð10:6Þ

since the concentration of a solid material is assigned as 1. When cation movements dominate the redox processes, the reaction is given as: PPyþ A þ X þ þ e ! PPyþ A X þ CPPyþ A X þ ¼ E0 þ 0:059 log CX þ ) E ¼ E0 0:059 log CPPyþ A
CX þ

ð10:7Þ

The exact positions of the potential peaks are derived from the differences between the electrode potential, E, and reference electrode potential, which is related to the reactions occurring at the reference electrode. Even so, Equations (10.6) and (10.7) clearly show that the redox potential depends on the dopant and electrolyte concentrations. Generally, in the case of anion movement, the redox potentials tend to decrease as the concentrations of these anions in the surrounding electrolyte increases. In the case of cation movement, however, the redox potentials are likely to increase with the increasing cation concentration. As mentioned above, the addition/removal of charge by electrochemically changing of conjugated polymer redox states and the flux of ions to balance the charge are the fundamentals for conjugated polymer actuators. Thus, the rate of actuation is primarily affected by the rate of charge transport, which is restricted by the internal resistance of the

204

Biomedical Applications of Electroactive Polymer Actuators

electrochemical cell and by the rate at which ions are transferred within the polymer [29, 30]. In general, charge transport is affected by the following factors: Polymer conductivity: Conductivity affects the current passing through the polymer especially when there are long electronic current paths. Low conductivity causes high potential drop along the polymer, so that the electrochemical potential is reduced leading to lower redox currents. Applied potential: The polymer resistance can be compensated by increasing applied potential in a controlled manner, resulting in improvement of charge transport. (However, irreversible degradation can also occur if very high electrochemical potentials are applied to the ICP.) Electrolyte conductivity: The concentration of ions in the electrolyte determines the conductivity of the electrolyte. Higher electrolyte conductivity will result in faster charging of the double layer and will also decrease resistive losses thus increasing the efficiency of charge transport. Ion diffusion: Ion diffusion rate is determined by the ion transport environment and the nature of the ions. At a given condition, smaller ions usually possess higher mobility than larger ions and can move more quickly into the polymer. Ion mobilities have also been shown to be dependent upon the prior swelling of the polymer, since conformational relaxation of the polymer can take some time to occur [31]. Polymer morphology and geometry: Polymer morphology depends on the conditions of electrochemical deposition. Usually porous structures of polymer will make it easy for ions to penetrate. The polymer geometry can also affect the charge rate. For example, if the same potential is applied along two polymer films in different thickness, the charge density increases faster in the thinner film than in the thicker film thus results in a faster actuation in the thinner film. The redox processes that drive actuation in ICPs are very much affected by ion transport processes, which are reviewed in more detail below.

10.5 Effect of Composition, Geometry and Electrolyte on Actuation of PPy 10.5.1

Effect of the Dopant Ion

One of the most studied influences on the actuation of PPy is the type of dopant ion used. The dopant ion is fixed during synthesis, but may be exchanged during subsequent redox cycling in a different electrolyte. The choice of dopant ion is almost endless, from ‘simple’ ions like chloride and nitrate to larger organic ions such as p-toluene sulfonate, dodecyl sulfonate or poly(styrene sulfonate). The dopant used during polymerization profoundly affects the morphology and density of the film formed at the electrode. Since actuation testing requires a mechanically robust polymer (especially when free-standing films are used), not all dopant ions are suitable. PPyþNO3 , for example, forms a powdery deposit and cannot be used to make films. Since the size of the dopant ion can be varied considerably, it is expected that the extent and speed of actuation will be very dependent on dopant ion size. The effect of dopant ion on actuation has been most commonly investigated by altering the composition of the test

Conjugated Polymer Actuators: Fundamentals

205

electrolyte after the polymer has been formed. Thus, the polymer is polymerized using a particular dopant ion that gives the desired structure and properties, but actuation tests are then carried out in a different electrolyte. In such cases there can be an exchange of dopant ion during the first several redox cycles. The initial dopant ion can be expelled during reduction of the polymer, but subsequent re-oxidation is likely to imbibe the alternative electrolyte anion, which is more likely present in larger quantities. The situation is simpler in cases where cation movement occurs, since no dopant exchange occurs. In one recent study, electrolyte composition was systematically varied to determine the effect of anion size on the actuation of an anion-dominated PPy actuator. Hara et al. [32] investigated the actuation characteristics of the anion induced movement of PPy doped with bis(trifluoromethanesulfonyl)imide (TFSI). The films were actuated in different electrolytes containing a homologous series of anions of increasing size with formula Li(CnF2nþ1SO2)2N, where n ¼ 1 (trifluoro: LiTFSI), n ¼ 2 (pentafluoro: LiPFSI) and n ¼ 4 (nonafluoro: LiNFSI). These workers found that the actuation speed increased as the size of the anion in solution decreased due to the fact that smaller ions have the ability to diffuse through the bulk of the polymer at a much faster rate. However, they also found that with increases in anion size, larger maximum strains were achievable. A summary of these key results may be seen in Table 10.2. Table 10.2 Actuation strain obtained from PPy.TFSI films tested in various electrolytes under isotonic conditions Electrolyte

Maximum strain (%) Contraction (%) After 2 s After 5 s After 10 s After 100 s

LiTFSI

LiPFSI

LiNFSI

24.1

26.1

30.0

5.9 11.2 15.1 20.1

5.3 11.1 16.1 22.9

3.0 7.3 12.9 28.0

Other studies have shown that the effect of dopant ion has a more complicated effect on the actuation process in polypyrrole. At some point with increasing anion size, the mobile ion exchanged between the polymer and the electrolyte changes from the anion to the cation. This effect has been clearly demonstrated by synthesizing polymers with different dopants. In one study [33], for example, PPy was electrochemically synthesized using p-toluene sulfonate and separately with the polymeric dopant poly(vinyl sulfonate) (PVS). When tested in various aqueous electrolytes, the PPy.PVS actuation was dominated by cation movements with the polymer expanding in the reduced state and contracting in the oxidized state. The PPy.pTS, however, showed more complex behaviour, as illustrated in Figure 10.6. Mixed ion movements occurred with reduction from the fully oxidized state first involving anion expulsion followed by cation incorporation as the polymer was further reduced. The two processes were reversed during the oxidation cycle. A further phenomenon known as ‘salt draining’ can also occur if the polymer is held in the reduced state [34]. Here cation / anion pairs can diffuse out of the neutral polymer causing slow shrinkage.

Biomedical Applications of Electroactive Polymer Actuators 1.5

0.4

2.0

1.0

0.4 Force (N)

0.3

0.5

Force (N)

Force (N)

0.5

Extension (mm)

1

1500

1600

1700

1800

1900

1.0 Force (N) Extension (mm)

0.1 1800 1900 2000 2100 2200 2300 2400

0.0 2000

3.0

3

(1.0) Potential (V)

(3.0)

1700 1800 Time (s)

(a)

0.5 0.0

0.5

0.0

1 –1 –3

(0.5) Current (mA) Potential (V)

Current (mA)

1600

Current (mA)

5 Current (mA)

Potential (V)

5.0

1.0

1500

1.5

Time (s)

0

–1 1400

4

0.2

Time (s) 0.5

–0.5

3

0.3

Extension (mm) 0.2 1400

2

Extension (mm)

0.6

Potential (V)

206

1900

(5.0) 2000

–5 1800 1900 2000 2100 2200 2300 2400

(1.0)

Time (s)

(b)

Figure 10.6 Actuation force generated under isometric conditions for PPy.pTS actuated in (a) sodium poly(vinyl sulfate) and (b) sodium nitrate electrolytes. Labels in (b) refer to anion movements (2 and 3) and cation movements (1 and 4) (Reproduced with permission from Synthetic Materials, Mechanism of electromechanical actuation of polpyrrole by Gandhi, M.R., Murray, P., Spinks, G.M. and Wallace, G.G., 73, 3, 247–56. Copyright (1995) Elsevier).

The extent of anion vs cation movement in the PPy.pTS was also found to depend on the electrolyte used. Monovalent cations (Naþ, Kþ) were found to be incorporated into the polymer during reduction to a greater degree than divalent cations (Mg2þ, Ca2þ) [33]. Okamoto et al. [35] also reported that in the cation derived movement of PPy.pTS benders in different aqueous solutions, there were significant differences in the extent and speed of actuation that was related to the size of the cation in solution. They found that the actuator moved more slowly with larger cations and that, over a certain cation size, neither the oxidation / reduction current nor bending was observed. 10.5.2

Effect of Solvent

Different solvents used in the electrolyte have been shown to have an impact on the extent of actuation achievable and this may attributed to two main factors:

 Electrolytic salts will have a different ionic conductivity in different solvents.  The extent of swelling and softening of the bulk polymer material will be different for different solvents. An improvement in the ionic conductivity will enable a larger and faster build up of charge in the polymeric material, since resistive losses are reduced. Higher electrolyte conductivities effectively mean that a higher electrochemical potential is experienced by the polymer working electrode for a given applied external voltage. The higher electrochemical potential can cause a greater change in the oxidation state of the polymer

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resulting in increased actuation. Likewise, solvents that have the ability to diffuse between polymer chains can enable both a plasticization of the polymer and an increase in the free volume. The more open polymer structure enables faster diffusion of ions within the polymer and the softening of the polymer can result in larger strains. These two effects were clearly seen in a study undertaken by Hara et al. [36]. In their work, PPy doped with bis(trifluoromethanesulfonyl)imide (TFSI) was actuated in various water/propylene carbonate (PC) solutions containing LiTFSI. The optimum performance of 23.6 % maximum strain at a strain rate of 10.8 % s 1 was achieved within an actuation solution that consisted of 60 % water and 40 % PC. Improvements in both the strain rate and the maximum strain were seen with actuation in LiTFSI electrolytes of water/PC blended solvents over actuation in electrolytes of either water or PC alone. The improved actuation was attributed to the fact that a greater swelling occurred from the PC solvent (enabling a faster and easier ion transfer) and an improvement in the ionic conductivity from the water solvent (enabling a better charge transfer). As such, the optimised performance for this system was realised at 40 % PC. Exchange of solvent between the polymer and the electrolyte has also been identified as a major contribution to actuation in PPy [18, 37]. The solvent exchange results from osmotic pressure differences that arise as a result of the different ion concentration inside and outside the polymer. The ion concentration within a highly doped PPy can be as high as 5 M [18]. When in contact with an electrolyte of lower concentration, solvent will diffuse into the polymer due to the osmotic pressure to try and even out the concentration gradient. This diffusion of solvent into the polymer causes the polymer to expand, adding to the actuation process. The reverse process occurs when the ion concentration inside the polymer decreases (e.g. during reduction of an anion-dominated PPy) so that solvent drains from the polymer causing further shrinkage. Ionic liquids are a special type of solvent-less electrolyte. Clearly no osmotic effects occur when ionic liquids are used for actuation studies. The complexity of ICP actuation is highlighted in PPy actuation studies in ionic liquid electrolytes compared with conventional solvent/dissolved ion electrolytes. In one study [38] using 1-butyl-3-methylimidazolium (BMI) hexafluorophosphate (PF6) or 1-ethyl-3-methylimidazolium (EMI) bis(trifluoromethanesulfonyl)amide (TFSA) ionic liquids the actuation was dominated by cation movements giving isotonic actuation strains of  3 % (CV scan at 50 mV/s between þ2.5 V and –2.5 V in a two-electrode cell). The same polymer tested in tetrabutylammonium.PF6 in propylene carbonate showed an anion-dominated actuation process. Strains of about 1.5 % were obtained under the same isotonic conditions. Ions form clusters in ionic liquids, so actuation is likely due to the diffusion of large positively charged clusters, rather than individual cations. The clusters break up when diluted with solvent, so actuation in the PC–PF6 electrolyte involves diffusion of individual ions plus associated solvent molecules and the osmotic movement of solvent. The ‘direction’ of actuation, then, is determined by the size and diffusion speed of these larger ionic assemblies. Even in conventional electrolytes, the type of solvent used influences the mobile ionic species. In PPy doped with ClO4 and actuated in LiClO4 in either propylene carbonate or acetonitrile, it was found that the anion (ClO4) was mobile in propylene carbonate, while in acetonitrile the cation (Liþ) was the mobile species [28]. A major advantage obtained from ionic liquid electrolytes, however, is a dramatically increased actuation stability on extended cycling. This effect was clearly illustrated in a

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Biomedical Applications of Electroactive Polymer Actuators 1.5

100 80

1.0 60 1 Hour “Rest Period” 0.5 PC-TBAPF6 0.0

0

2000 4000 Cycle Number

40 20

6000

Norm. redox charge (%)

Linear displacement (%)

[BMIM][PF6]

0

Figure 10.7 Actuator strain measured during repeated cycles at 1 Hz using a potential pulse of þ5 V and –5 V in either a conventional (TBA.PF6 in propylene carbonate) electrolyte or an ionic liquid (BMIM.PF6). The charge passed per cycle is also plotted and the changes in actuation upon repeated cycling is related to the decrease in charge passed (Reprinted from Science, Use of Ionic Liquids for – Conjugated Polymer Electrochemical Devices by Lu, W., Fadeev, A.G., Qi, B., et al. with permission, 297, 5583, 983–7. Copyright (2002) AAAS).

long-term cycling study of PPy actuators in the ‘conventional’ TBA.PF6 in PC electrolyte compared with an ionic liquid (BMI.PF6) [38, 39]. The actuation strain slowly and steadily decreased upon repeated cycling in the conventional electrolyte (Figure 10.7). A much smaller decline in the actuation strain was observed in the ionic liquid electrolyte. In both cases the actuation strain correlated with the amount of redox charge passed during the redox cycle. The amount of charge declined steadily in the PC electrolyte, but remained quite stable in the ionic liquid. These behaviours were attributed to a slow degradation of the PPy in the PC electrolyte associated with parasitic redox reactions involving the solvent. Such reactions were absent in the ionic liquid (IL), so the PPy remained undegraded and maintained a reasonably consistent actuation over several thousand cycles. The stability of actuators is also influenced by the mechanical load applied and a further benefit of IL electrolytes is highlighted by such studies [38]. As shown in Figure 10.8 the actuation strain decreases rapidly in a PC electrolyte (TBA.PF6) when actuation strain was measured at higher isotonic stresses. The strain remains approximately the same, however, as the stress increased and an IL electrolyte was used. The effect of applied stress on the actuation strain observed is related to the Young’s modulus of the polymer, as explained below. 10.5.3

Charge Transfer Processes

Many of the factors that have been discussed previously relate to the ability of the ions in the electrolyte solution to pass into or out of the free volume of the polymer and cause a resulting expansion or contraction. However, this movement is driven by a transfer of charge on the polymer backbone, and if this transfer of charge is restricted, then so too is the extent of actuation. The most important factor involved with electronic charge transfer is

Conjugated Polymer Actuators: Fundamentals

209

3.5

Strain (%)

3 2.5 2

IL

1.5

PC

1 0.5 0 0

1

2

3

4

5

6

7

Stress (MPa)

Figure 10.8 Actuator strain measured at different applied isotonic stresses using a potential scan between þ2.5 V and –2.5 V at 50 mV/s in either a conventional (TBA.PF6 in propylene carbonate) electrolyte or an ionic liquid (IL: EMI.TFSA) (Reprinted with permission from Chemistry of Materials, Use of Ionic Liquids as Electrolytes in Electromechanical Actuator Systems Based on Inherently Conducting Polymers by Ding, J., Zhou, D., Spinks, G.M., et al., 15, 12, 2392–8. Copyright (2003) American Chemical Society).

the conductivity of the polymer. If the polymeric resistance is high, then electron flow through the polymer is restricted and it will be more difficult to switch the oxidation state of the polymer. Similarly, ion movements within the polymer must accompany the change in oxidation state of the polymer. Thus, the rate and extent of actuation will be affected by ion diffusion kinetics. Factors that are known to affect the resulting conductivity of the polymer include:

 type of dopant used  electron flow rate during polymerization  polymerization solution temperature. A thorough analysis of these and other factors is beyond the scope of this chapter and has been covered elsewhere [40, 41]. The rate of electronic charge injection is particularly important when fast actuation is desired, or when large actuators are to be used. Because of the poor conductivity of ICPs, there will be considerable IR (current  resistance) losses at positions some distance from the point of electrical connection. The film resistance means that the actual potential applied to the polymer decreases with distance from the connection point. Thus, full actuation may only occur near the connection point [42]. Ding et al. showed significant improvements in the speed of actuation of PPy tube devices by the incorporation of a thin helical wire of platinum, encapsulated within the device during growth [17], as shown in Figure 10.9. The platinum wire enabled better transfer of electrons through the length of the tube, ensuring that the bulk of the PPy material was able to be switched in oxidation state, together, and limited the time-lag effect associated with the lowered conductivity that is seen when switched to the reduced state. Similar structures have been built using tungsten springs [43]. Resistive losses also occur within the electrolyte, and Madden et al. have shown that resistance compensation techniques increase the rate of actuation of PPy films [44].

210

Biomedical Applications of Electroactive Polymer Actuators 1 0.8

(c)

Strain (%)

0.6 0.4 0.2

(b)

0 –0.2

(a) 8

12

16

20

Time (s)

Figure 10.9 Polypyrrole ‘helix tube’ actuator incorporating a thin platinum wire within the wall of the hollow tube and improved strain response of (a) a flat PPy film, ( b) hollow PPy tube and (c) the helix tube (Reproduced with permission from Synthetic Metals, High performance conducting polymer actuators utilising geometry and helical wire interconnects by Ding, J., Liu, L., Spinks, G.M., et al., 138, 3, 391–8. Copyright (2003) Elsevier).

As described above, the build up of charge on the polymer backbone and the incorporation of ions to balance the charge occur simultaneously. One process cannot happen without the other. Ion diffusion can sometimes be the rate limiting process in actuation and ion diffusion is directly related to the volume changes that drive actuation. Smela and co-workers [45, 46, 47] have carefully studied the mechanisms of ion diffusion in PPy using thin films sandwiched between a conductive layer and a transparent ion barrier layer. In this configuration, ion diffusion can only occur from the edges and the progress of diffusion can be monitored by both colour change and film thickness (in the latter case an elastomeric ion barrier was used). These studies used PPy.DBS, which is known to involve cation transport and show that:

 The first reduction cycle was different to later cycles. This effect has also been commonly observed in electrochemical studies on ICPs and is sometimes called the ‘breakin effect’ [48]. It is believed that the first swelling cycle is not completely reversed during subsequent shrinking. Although the swelling ions (Naþ) were removed during the first oxidation of PPy.DBS (in aqueous Na.DBS electrolyte), the free volume left behind is not completely removed by polymer chain relaxation [34]. This relaxation occurs by chains adopting a more compact conformation and takes some time to occur. Thus, second and later reduction cycles occur more easily (at less negative potentials) with 10 times faster ion diffusion rates.  Ions diffuse into the polymer with a distinct boundary between a fully reduced/ swollen phase and a fully oxidized/contracted phase. The boundary is very sharp on the first reduction cycle but is broader on subsequent cycles. This type of diffusion behaviour is typical of solvent diffusion in polymers that causes a softening phase change and is called Type II (non-Fickian) behaviour [49]. This process occurs in

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glassy polymers where the diffusing solvent molecules plasticizes the polymer, reducing the glass transition temperature to below the test temperature so that the solvent filled polymer becomes rubbery. Diffusion is much faster in the open rubbery phase than the compact glassy phase, so that the rate limiting step in diffusion is the plasticization of the polymer. A sharp boundary then forms between the solvent filled rubbery phase and the solvent-less glassy phase. Clearly, similar processes occur during the swelling of PPy during redox cycles. The differences between the first and subsequent redox cycles may again reflect the ‘opening-up’ of the structure that occurs during the first reduction.  Oxidation processes have not yet been considered, since they are complicated by polymer conformational relaxation processes and the low electronic conductivity of the starting fully reduced state. Other workers have considered ionic diffusion through the thickness of a PPy.ClO4 film during oxidation [50], but these results cannot be directly compared with Smela et al. because in one case the anion is dominant and in the other case the cation is dominant.  A direct correlation exists between the redox state and the PPy film volume. Although, not surprisingly, this correlation confirms that the ion diffusion drives the actuation process. The film thickness was measured by profilometry across the phase boundary. A clear height step was observed at the boundary between the oxidized and reduced phases. The thickness of polymer was constant in the reduced (swollen) region behind the boundary, showing that the polymer was fully reduced in this entire region. Thus, reduction of the polymer requires the diffusing ions to reach the next reduction site on the polymer.  The speed of reduction depends mostly on the applied voltage, with a small effect due to temperature [47]. The speed of the moving front increases approximately linearly for applied potentials below –0.8 V (vs Ag/AgCl). The reduction peak potential occurs at – 0.65 V in this system, so a significant overpotential is needed to drive the ion diffusion at reasonable speeds. Below –1.6 V the ion speed was constant, indicating that other processes became rate limiting under such conditions. Ion diffusion in the electrolyte or electronic conductivity through the film thickness may be the limiting processes at these potentials. Thicker films (to 2 mm) showed slightly faster moving front speeds, both in the first and subsequent cycles. The reasons for this effect are not known. Moving front speeds increased very slightly with increase in temperature from 20 to 40 oC. These studies highlight the key role of ion transport in determining the amount of actuation that occurs in ICPs and, particularly, highlights the slow nature of the ion transport. The absolute speed of ion transport determined in these studies was of the order of 20 mm/s. In most actuator devices, the diffusion direction is through the film thickness and it is not known whether the ion diffusion speeds are different in this direction compared with along the film length, as measured by Smela and co-workers. However, if we assume that the diffusion speeds are appropriate for the thickness direction, then a film 20 mm thick can only be fully reduced in 1 s. Thus, operation above 1 Hz would see a drop off in the extent of actuation since the film cannot be fully reduced/oxidized in the time available. Studies on the electrochemical efficiency of PPy films switched at 1 Hz showed a significant decrease in the amount of charge transferred for films greater than 0.5 mm thick [17]. Actuation studies on PPy bending actuators operated in air show a decrease in bending amplitude above 1 Hz [51]. These results all support an ion-diffusion limited

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charging and, hence, actuation in PPy. Other effects will also influence the speed of actuation, in particular the insulating nature of the reduced polymer will limit the speed of electronic charge injection. 10.5.4

Effect of Porosity/Morphology

The preceding discussion highlights that the free volume and pore size that are generated during growth of the polymer are critical to the performance of the actuator. It can be seen that there are various factors that can affect the morphology and pore size (or free volume) of the polymer material. Electrochemical growth on different substrates (such as gold, indium tin-oxide (ITO), glassy carbon (GC) and platinum) [52] has been shown to generate structures that have a varied pore size, morphology and actuation characteristics. Hara et al. evaluated that with films grown on these electrodes the actuator performance (max strain) followed that: GC > Pt > Au > ITO. SEM photographs of the corresponding films had similar microstructures on the solution side of the electrode. However, on the electrode side vast differences were evident. GC was seen to have an even distribution of fine pores approximately 1–2 mm in diameter, gold had a varied distribution ranging from 1–15 mm, platinum had an even distribution of pores approximately 8 mm in diameter and ITO had many fewer pores ranging from 5–15 mm in diameter. The anions within solution have ionic radii and solvation spheres in the nanometre range, and sub-micron sized structures that are more likely to be related to direct expansion were not visible under this magnification and may not be directly commented on. However, it was evident from the results that the strain rate of each of the polymers followed that: Pt > Au > ITO > GC. It is expected that the high prevalence of large pores in the polymer grown on platinum or gold enables easier passage of electrolyte into the bulk of the polymer, thus improving the rate of diffusion of ions into the sub-micron sized free volume of the polymer and improving the actuation strain rate. An earlier study by Pandey et al. [53] showed that the polymerization electrode also influenced the balance of anion/cation movement in the polymer. In this case PPy doped with naphthalene sulfonic acid (NSA) was prepared on three different electrodes. Anion movement was favoured in those films that had a more open, porous structure. The actuation was performed in aqueous NaCl electrolyte, so that the mobile cation (Naþ) was smaller than the mobile anion (NSA). As well as the electrode material, there are other factors that may help control the morphology of the deposited polymer. Electropolymerization by different means (cyclic voltammetry (CV), constant applied potential and constant applied current) are factors that are also known to affect the morphology of the deposited polymer. In general, constant applied potential generates films with a cauliflower structure that have a large free volume in contrast to the compact flat films that may be generated with growth by CV or constant current. By lowering the temperature or viscosity of the solvent, more compact films may also be generated that have improved conductivities and smaller pore sizes.

10.6 Mechanical System Response Ultimately, the actuation behaviour is determined by the mechanical output as manifest by a movement and/or a force. In practical actuator devices the required mechanical output

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213

will be variable – the force applied by the external system will vary and the required forces/ movements will change depending upon the desired outcome. The ICP’s mechanical properties (especially the Young’s modulus) will, therefore, have a bearing on the actuation response. It is now well established that the Young’s modulus of ICPs depends on the oxidation state [19, 54], so that the modulus will change during the actuation cycle. A simple model is described that links the length change that occurs with no external forces acting on the ICP to the displacement produced with an applied load [19, 55]. The importance of the Young’s modulus of the actuator material is demonstrated by a simple example. As illustrated in Figure 10.10a., a common device geometry has the actuator attached to a restoring spring. This geometry is especially useful when the actuator

Spring

Final State

f=0

f>0

l = l0 E=Y

Actuator

l = lf E = Y’

Actuator Displacement (mm)

Initial State

1.2 1 0.8

B

0.6

C

0.4 0.2 0 0.01

0.1

100

70 Work per cycle (kJ/m^3)

2.5 Force Generated (N)

1 10 Stiffness Ratio

(b)

(a)

2 B

1.5 1

A

0.5 0 0.01

A

C

0.1

1 10 Stiffness Ratio

(c)

100

B

60 50 40 30

A

20 C

10 0 0.01

0.1

1 10 Stiffness Ratio

100

(d)

Figure 10.10 (a) A contractile linear actuator attached in series with a restoring spring; (b) final contraction (actuator stroke) achieved when operated against springs of increasing stiffness (stiffness ratio is the ratio of the spring stiffness to the actuator materials stiffness) and different Young’s moduli: Y is the Young’s modulus in the expanded (initial) state and Y’ is the Young’s modulus in the contracted (final) state; (c) force generated; and (d) work per cycle. In (b), (c) and (d) the labels for the curves represent: A: Y ¼ Y’; B: Y ¼ 1/2 Y’; C: Y ¼ 2Y’.

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Biomedical Applications of Electroactive Polymer Actuators

is long and slender, as in a film or fibre. In such geometries the actuator does not ‘push’ well due to the ease of buckling. To overcome this problem, the actuator is configured to operate in contraction and the restoring spring ensures expansion in the reverse cycle. The actuator/spring model is also approximately correct for the bender actuator, since the passive layer must be bent as the actuator contracts/expands and the bending of the passive layer produces a restoring force. In situations where the external spring force (fe) depends linearly on its displacement with a spring constant (ke), it has been shown [55] that the final change in length of the actuator will be:   1 DLf ¼ DL0 ð10:8Þ 1 þ r0 where DL0 is the actuation displacement when no external force is applied (the ‘free stroke’) and r’ is the ratio of the external spring stiffness (ke) to the stiffness of the actuator material (ki’). The internal stiffness is of the actuator material in its final state and takes into account any changes in Young’s modulus (Y’) that occurs during actuation: r0 ¼

ke k0i

and

k0i ¼

Y 0A L0

ð10:9Þ

where A and L0 are the initial cross-sectional area and length of the actuator, respectively. As illustrated in Figure 10.10b, the amount of actuation achieved depends significantly on the stiffness of the external spring. Low stiffness ratios correspond to a soft external spring, which then applies only a small force to the actuator. The actuator then produces close to its maximum displacement (DLf » DL0 ), since little elastic stretching by the spring occurs. When a stiff external spring is used (high r’), however, there can be little or no actuation displacement (DLf » 0), since the high force applied by small displacements of the spring cause elastic stretching of the actuator that counter balance the actuation contraction. The force generated by the actuator (Figure 10.10c) shows the opposite trend: low forces when a soft spring is used and maximum forces when a stiff spring is used. The work performed by the actuator in deforming the spring is given in Figure 10.10d and reaches a maximum when the external and internal stiffness ratios are equal (r’ ¼ 1). The data presented in Figure 10.10 are for a hypothetical actuator that generates a contraction of 1 mm, has original dimensions of L0 ¼ 20 mm and A ¼ 0.2 mm2 and initial Young’s moduli of Y ¼ 100 MPa. Also shown in Figure 10.10 is the effect of changing modulus of the actuator material on the actuation performance. The stiffness ratio on the x-axis is calculated assuming no change in modulus during actuation (Y’ ¼ Y ). The actuation displacement, force generated and work per cycle have all been calculated for three conditions: Y’ ¼ Y; Y’ ¼ 0.5Y and Y’ ¼ 2Y. It can be seen that for a given external spring stiffness, the actuator displacement decreases when the modulus decreases during the contractile actuation. The lower actuator modulus means that the material will stretch to a greater extent due to the external force applied by the spring. The net contraction is, therefore, smaller. Similarly, the force generated is smaller as a result of the smaller net contraction. In contrast, the actuator displacement, force generated and work per cycle all increase when the modulus increases during contractile actuation.

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215

The calculated results given in Figure 10.8 all assume that the free stroke of the polymer does not change in each circumstance considered. Although experimental data does support the findings illustrated in Figure 8, it has generally been found that situations that change the modulus of the actuator also change the free stroke. For example, it is shown from Equation (10.8) that the actuation displacement would increase if the modulus in the contracted state was increased (and the free stroke unchanged). Addition of carbon nanotubes to polyaniline fibres certainly increased the modulus [56], however, the presence of the nanotubes also restricted the actuation of the PANi so that the free stroke also decreased. The result shown in Figure 10.11 was that the reinforced PANi gave higher actuations when higher external forces were applied, but the unreinforced PANi gave the larger actuation at smaller external forces [56]. These data were collected under isotonic conditions (rather than using a restoring spring), so that the actuation displacement is now given by Equation (10.10):   f L0 1 1 DLf ¼ DL0 þ ð10:10Þ A Y0 Y where f is the applied isotonic force. The graph shown in Figure 10.11 approximates the linear behaviour predicted in Equation (10.10), with the intercept corresponding to the free stroke and the slope determined by the modulus shift during actuation. 0.2 0

Strain [%]

–0.2 –0.4 –0.6 –0.8 –1 –1.2 –1.4

0

25

50 75 Stress [MPa]

100

125

Figure 10.11 Actuation strain obtained under isotonic conditions for PANi fibre (square symbols) and carbon-nanotube reinforced PANi fibre (circles). A negative strain indicates contraction.

A number of studies on different ICPs have shown similar data to that given in Figure 10.11, with a decreasing actuation occurring when higher isotonic stresses are applied [19, 24, 57]. These observations reflect the fact that, in most cases, the modulus is smaller in the contracted state. The modulus of solvent swollen network polymers is known to be influenced by two factors. Firstly, the swelling by solvent reduces the concentration of load-bearing chains, so that the modulus tends to decrease. Secondly, as

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Biomedical Applications of Electroactive Polymer Actuators

chains become highly extended as a result of solvent swelling, they effectively become stiffer, which leads to an increase in modulus. This latter effect appears to dominate in most ICPs. The dependence of modulus on swelling is complex, however [58], and not linearly dependent upon the degree of swelling. It has been shown, for example, that there is almost no modulus change when a PPy material was subject to a small voltage range [19]. The same material, however, showed very large changes in modulus when a wide potential range was used to cause full oxidation and reduction of the polymer [19]. In this case the modulus changed by a factor of four, causing a very sharp drop in actuator strain as the applied stress increased. There has been at least one report of an increase in actuator strain with increasing stress. This situation arises when the modulus increases in the contracted state and occurred in a PTh film tested in an ionic liquid electrolyte [59]. Curiously, the same polymer showed the more typical ‘low modulus in the contracted state’ when tested in a conventional electrolyte, highlighting the complexity of the modulus shift phenomenon. Other mechanical effects also influence the actuation behaviour of ICPs. An unresolved problem at present is the creep or drift that is commonly observed. An example is shown in Figure 10.12, with cyclic potentials producing a larger expansion per cycle than the corresponding contraction [60]. The result is a net increasing in length of the sample over time. The rate of this drift is increased when higher loads are applied to the actuator, implying that the drift is related to a viscous or viscoelastic deformation occurring in the

20 Isotonic load (MPa) 34

Strain (%) curves offset for clarity

16

28 23

12

17 11

8

8.5 5.7

4

a c b

3.4 1.1

0

1.3% strain 0

100

200 300 Time (s)

400

500

Figure 10.12 Example data showing drift in actuation strain over time for polyaniline actuated at different applied isotonic stresses (Reproduced with permission from Synthetic Materials, Polyalinine actuators: Part 1. PANI(AMPS) in HCI by Smela, E., Lu, W. and Mattes, B.R., 151, 1, 25–42. Copyright (2005) Elsevier).

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217

polymer. In the case of polyaniline, a significant decrease in the rate of drift has been achieved by adding carbon nanotube reinforcement [56]. However, as described above these reinforcing agents also reduce the amount of actuation occurring. More research is needed into the origin of the drift and the molecular processes occurring, before the problem can be rectified.

10.7 Device Design and Optimization The size and shape of the polymer actuators are as important as their electrochemomechanical properties when it comes to providing enough actuation power for practical applications. The actuation power depends on (i) force output, (ii) displacement output and (iii) speed of response. In this section, a practical approach is presented to link device design requirements to the performance parameters and to offer guidelines for the device design and optimization based on electroactive PPy bending-type actuators. The actuator considered is a one-end fixed and the other end-free bender. 10.7.1

How to Tailor Actuator Performance to Meet Design Requirements

The force and displacement outputs are created due to internal bending moment induced during the conversion of electrochemical energy into mechanical energy. The force (blocking force), F, can be estimated using a quantitative relationship based on the strain created as a function of the input voltage [61]: F¼

E1 b  h1 ðh1 þ h2 Þ L

ð10:11Þ

where is the Young’s modulus of the PPy layers. For other variables in Equation (10.11), see Figure 10.13. The strain a in the PPy layers is a function of the strain to charge ratio and charge density in the PPy layers [62, 61]. As described in Equation (10.11), the force output h1

h2

Upper PPy layer

dx h1

y

Neutral axis Lower PPy layer dθ

R

F Figure 10.13 One end cantilevered actuator and the model parameters (Reproduced with permission from Bioinspir. Biomim., Establishment of a biomimetic device based on tri-layer polymer actuators-propulsion fins by Alici, G., et al., 2, S18–30. Copyright (2007) IOP).

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Biomedical Applications of Electroactive Polymer Actuators

Input Power (W)

Max Force (mN)

14 12 10 8 6 4 2 0 9

6

5

4

3

0.016 0.014 0.012 0.01 0.008 0.006 0.004 0.002 0

2

9

Actuator Width (mm)

6

5

4

3

2

Actuator Width (mm)

Figure 10.14 Variation of the maximum blocking force and the input power with the actuator width. The actuators used were 15 mm long with 30 m PPy thicknesses under 1 V. The electropolymerization of polypyrrole is achieved by submerging the sputter coated PVDF film in a solution of 0.1 M pyrrole, 0.1 M LiTFSI in Propylene Carbonate (PC) with 0.5 w/w% water (Reproduced from Electroactive Polymer Actuators and Devices (EAPAD) 2007, Proceedings of SPIE Vol. 6524, Tri-layer conducting polymer actuators with variable dimensions by Minato, R., Alici, G., McGovern, S. and Spinks, G., 6524, 6524J. Copyright (2007) SPIE).

is proportional to the width, and thickness, and inversely proportional to the actuator length. This follows that appropriately sized actuators can satisfy the force design requirement. For the force results presented in Figure 10.14, the force output increases with the actuator width. For the widths >4 mm, the actuator curls into a semi-cylindrical shape. This curling action decreases the bending displacement. However, the time-averaged electric power input increases with the actuator width [63]. The experimental bending moment (F  L) of the PPy actuators as a function of PPy thickness is provided in Figure 10.15, where there is a non-negligible drop in the bending

Bending Moment (mN.cm)

1.4 1.2 Experimental

1

Theoretical 0.8 0.6 0.4 0.2 0 0

30

60 90 Thickness (µm)

120

Figure 10.15 Experimental and theoretical results showing the variation in the bending moment with the thickness for an actuator with the dimensions of 20  1  0.17 mm and for the salt TBA.PF6 0.25 M in the solvent propylene carbonate.

Conjugated Polymer Actuators: Fundamentals

219

Bending Displacement (mm)

moment output when each of the PPy layers is thicker than 60 mm [64]. When the PPy layer is thicker than 60 mm, the bending moment begins to decrease. In this case, the PPy actuator is unable to overcome its increased flexural rigidity. For instance, an increase in the PPy thickness from 30 to 50 mm gives a 25 % increase to the total thickness of the strip, that is a 66 % increase in the volume of the PPy, but the flexural rigidity increases by 95 %. The corresponding bending displacement shows a similar trend. 8 7 6 5 4 3 2 1 0

Vertical Horizontal

0

0.2

0.4 0.6 Voltage (V)

0.8

1.0

Figure 10.16 The variation of the displacements with constant input voltages for actuator with the dimensions of 10  1  0.17 mm for the electrolyte TBA.PF6 0.05 M in the solvent propylene carbonate.

As depicted in Figure 10.16, the bending displacement or the tip displacement of the actuators is proportional to the input voltage [61], and the following differential equation describes the horizontal ‘x’ and vertical ‘v’ displacements of the actuators: "  2 # d2 v dv 3 V C ðh2 þ h1 Þ "  1þ ð10:12Þ   3 # ¼ 0 2 dx dx 2h1 þ h2 3 h2 þ ð E2 E 1 Þ 2 b L E1 2 2 where V is the input voltage, L is the length of the actuator, b is the width of the actuator, C is the capacitance, E1 and E2 are Young moduli of PPy and PVDF layers, respectively, h1 and h2 are described in Figure 10.13, and is the experimentally determined proportionality constant relating the internal stress s to the exchanged charge density. The experimental and theoretical results provided here suggest that, depending on the force, displacement and speed requirements of a practical device, the geometric parameters and shape of the actuator can be optimized suitably to satisfy the device requirements. As a case study, a swimming device propelled with the bending type actuators is presented next to demonstrate the influence of the actuator geometry on a functional system. 10.7.2

Design of a Swimming Device

As reported before [65, 66, 67, 68, 69], when designing a swimming device, there are two paramount factors to consider: (i) the shape, mechanical properties and the locations of propulsors/fins on the device; and (ii) their movement pattern. Fish is a good example to

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follow in considering these factors appropriately. The propulsion efficiency or the swimming efficiency is governed by, the Froude number: ¼

TV P

ð10:13Þ

where T, V and P are the thrust generated, the velocity of the swimming body and power input. T and V are determined by the three design parameters outlined in the opening paragraph of this section. The thrust generated by a swimming device depends on four factors [67]: (i) the aspect ratio (the higher is the aspect ratio, the higher is the net thrust force to accelerate the device); (ii) the shape of the fins; (iii) the fin stiffness (the higher is the stiffness, higher is the thrust); and (iv) the oscillatory motion of the fins (the higher are the frequency and amplitude of the fin oscillations, the higher is the thrust). With reference to the performance results of the polymer actuators presented above, all these factors can be controlled except the fin stiffness, which highly depends on mechano-electro-chemical properties of the actuators. The topology of the proposed device is depicted in Figure 10.17 where eight fins are installed along both sides of a rigid body to move the device in a direction perpendicular to the longitudinal axis of the body. The rigid body is made of prepregnated carbon fibre strips of 0.3 mm thickness and hardened with resin. The device can be considered like a box fish having a carapace (rigid body) with side or paired fins running through the rigid body, like a fish having pectoral fins. The fins or polymer benders can be considered as individually controlled flexible membranes. Each fin is activated with sinusoidal inputs such that there is a phase lag between the movements of the successive fins, and the frequency and amplitude of the input can be changed to create enough thrust for propulsion. This generates an undulatory movement. This is in agreement with the finding in the literature [69] that the undulatory movement is superior over an oscillatory movement, from the propulsion efficiency point of view.

Paired fins

chord span

Direction of movement

Figure 10.17 Topology of the proposed biomimetic device (not to scale) (Reproduced with permission from Bioinspir. Biomim., Establishment of a biomimetic device based on tri-layer polymer actuators-propulsion fins by Alici, G., et al., 2, S18–30. Copyright (2007) IOP).

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With reference to Figure 10.17, the aspect ratio, which is the ratio of the fin span to the fin chord, characterizes the movement of swimming bodies. A low aspect ratio fin generates a low velocity of movement, but results in great propulsion efficiency and manoeuvrability. As far as the force output of polymer actuators is concerned, low aspect ratio fins are not favourable. The biomimetic devices that are currently being worked on have high aspect ratios, which create less drag force. This follows that the devices with caudal fins are preferable. Further, future designs will also include fins at both ends of the body – carapace, like caudal fins. As reported in the biomimetic literature, such a device has better dynamic stability and endurance [67]. Depending on the direction of motion, either of both tail fins can be activated for cruising. 10.7.3

Device Testing

The assembled rigid frame and the fins are shown in Figure 10.18. Undulatory movements were employed to create enough thrust for propulsion [70]. This was achieved by activating each fin with a sinusoidal input such that there was a phase lag between the movement/operation of the successive fins. The best propulsion was observed at 1.5 Hz, 2 V peak to peak, with phase delays of 90o. A test was conducted when the first fin of one side of the prototype and the last fin of the other side were connected to the same power source, moving simultaneously. Similarly, the second fin of one side was connected to the second last fin from the other side, and so on. It was observed that each fin created sufficient undulatory movement to cause the prototype to rotate approximately 5°, though the platinum wires restricted the propulsion motion of undulating fins. This follows that the placement and method of attachment of the polymer actuator to the mechanical device determine the direction of movement and defines the actuator requirements.

Carbon Pre-preg Rigid Body Platinum Wire

Polymer Actuator Fins

Figure 10.18 The prototype swimming device with polymer actuators as the propulsion elements – fins (Reproduced with permission from Bioinspir. Biomim., Establishment of a biomimetic device based on tri-layer polymer actuators-propulsion fins by Alici, G., et al., 2, S18–30. Copyright (2007) IOP).

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The next test was conducted by connecting the first fin of one side to the first fin of the other side of the prototype. All the other fins were connected in the same manner. The propulsions of the fins were large enough to produce a side-to-side motion along the longitudinal axis of the fins. Although this experiment was successful in a small container, it was not possible to observe the full distance the prototype could move under the propulsion force produced by the fins. More importantly, this test demonstrated that the fins made of bending type polymer actuators could generate enough thrust to move the device in an aquatic medium. In summary, the performance quantification of bending polymer actuators have been presented in terms of the force and displacement outputs, and a methodology has subsequently been proposed to design a swimming device propelled by the polymer actuators, which act as the propulsion elements of the device. All performance results presented and the successful testing of the device demonstrate the importance of understanding the application requirements and then tailoring the geometry and shape of the actuators, and the placement and attachment of the actuators for a successful attempt to widen the application areas of ICPs.

10.8 Future Prospects Since the first bending-type actuators from ICPs were demonstrated in the early 1990s, there have been considerable advances in actuator performance, applications and in the understanding of actuation mechanisms. In terms of performance improvements, most attention has been given to free-standing films. Around the turn of the century, the best ICP actuation performances (reported by several different groups) were around 5 % maximum strain, 1 %/s maximum strain rate and 5 MPa maximum sustained stress (isotonic). Since 2000, significant improvements have been reported in each of these areas:

 maximum actuation strain: 12 % in 2003 [71]; 26 % in 2004 [72]; 40 % in 2005 [32]  maximum actuation strain rate: 4 % in 2000 [44]; 15 % in 2003 [17]  maximum sustained stress: 34 MPa in 2002 [73], 100 MPa in 2006 [56] While these improvements are impressive, the maximum performances in each area have not been achieved simultaneously. The highest stress actuators, for example, produce an actuation strain of only  2 % [56]. It is particularly useful to have actuators that give simultaneously high stroke, fast response and can operate against high stresses. A web based resource for tracking the published actuator performances of ICPs (and other actuator materials) has been developed by the Molecular Mechatronics Group at the University of British Columbia [74]. While the focus of much of the research to date has been in understanding mechanisms and improving the basic performance, several other areas require further work. Improvements can be achieved in the following areas:

 Efficiency: the energy conversion efficiency (electrical to mechanical) of ICP actuators is poor at <1 % [75]. Much of the electrical energy input is stored as Faradaic charge in the actuator material and possibly could be extracted to improve efficiency.

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 Creep/drift/stability: most ICP actuators suffer stability problems when stressed, which makes precise positioning difficult. Little is known about the mechanisms of the positional drift.  Modelling: the complexity of the actuation process means that a full description of the actuation performance is still lacking. A quantifiable model linking the electrical input (voltage/current) to the mechanical output (position/force/speed) is required for real applications.  Scale up: most ICP actuators have been produced in relatively thin films usually 10–20 mm in length and  5–10 mm wide. While the stresses and strains produced are impressive, the actual forces and displacements are small. Large scale applications (e.g. in humanoid robots) will require tens of millimetres in displacement and tens of Newtons in force. Parallel and serial assemblies of ICP actuators are required to deliver these performances. Considering the low energy conversion efficiency of ICPs, the electrical power supply for such large actuators will be a serious limitation.  Scale down: the microelectromechanical systems (MEMS) industry offers tremendous opportunities for novel actuators and ICPs are well suited to meet these needs, as reviewed recently [76]. Fabrication, design and control issues need to be refined to develop useful micro-actuator devices.

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11 Steerable Catheters Tina Shoa1, John D. Madden1, Nigel R. Munce2 and Victor X.D. Yang2, 3 1

Advanced Materials and Process Engineering Laboratory and Department of Electrical & Computer Engineering, University of British Columbia, Vancouver, Canada 2 Imaging Research, Sunnybrook Health Science Centre, University of Toronto, Canada 3 Department of Electrical and Computer Engineering, Ryerson University, Toronto, Canada

11.1 Introduction Conjugated (or conducting) polymer (CP) actuators are a type of electroactive polymer (EAP) actuator, with particular characteristics which are advantageous for minimally invasive surgical and diagnostic tools. Some of the characteristics include low actuation voltage, ease of fabrication, relatively high strain and biocompatibility. In this chapter the application of conjugated polymer actuators in steerable catheters is discussed. Although various active catheters driven by shaped memory alloys (SMA), piezoelectric materials and microelectromechanical systems (MEMS) based devices have been presented, no active catheters are in wide spread use. Conjugated polymer actuators have shown attractive properties, which make them promising to be employed extensively in active catheter applications. In this chapter, performance specifications and regulatory requirements are presented, and several steerable catheter prototypes are described. The promise and challenges associated with conjugated polymer driven devices are discussed.

11.2 Catheters: History and Current Applications A catheter is a thin flexible hollow tube that is inserted into body cavities to provide a channel for fluid passage or an entry for a medical device. Catheters as fluid channel may be used to remove waste fluids from the body or direct a liquid into the body. Practical

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applications include draining urine from the urinary bladder, injecting intravenous fluids and administrating medication or nutrition directly into the body. Catheters are also used to direct a medical device to a particular part of the body for minimally invasive diagnosis and treatment procedures. In angiography a catheter is used to administer the radio-contrast agent at the desired area to be visualized. In angioplasty and stenting, catheters are employed to guide a therapeutic device to open a blockage inside a vessel. During interventional neurovascular angiography, flexible coils can be delivered by extremely small catheters to treat cerebral aneurysms. Intravascular ultrasound (IVUS) is another example where a catheter based miniature ultrasound transducer is used to provide ultrasonic images from the inside of blood vessels. In Optical Coherence Tomography (OCT) an optical fibre is guided through a catheter to a particular part of the body to generate high resolution three-dimensional (3-D) images. Flexible endoscopes, used for visualizing natural body lumens, such as the airway or gastrointestinal tract, represent large catheters with imaging sensors which can guide diagnostic or therapeutic interventions inside a patient body such as biopsy, resection, ablation or injection. Catheters have a long history [1]. The earliest catheters were formed from straw and leaves and were used for drainage of urine. Hollow leaves of plants, coated with lacquer, were used as catheters in China around 100 BC. In 1036 AD Avicenna described the first flexible catheter made from stiffened animal skins. Benjamin Franklin designed a flexible silver catheter in 1752 for patients who suffered from bladder disease. The modern application of the catheter started in 1868 when a catheter with features for controlling the catheter insertion depth was patented by Dr N.B. Sornborger. The modern disposable catheter was invented in the 1940s by David S. Sheridan [1]. Flexible disposable catheters are used in many applications and their structures and designs have been improved over the last 60 years. The conventional method of handling A catheter involves inserting it into the body passively, by pushing it from outside. Guide wires, manipulated externally from the patient, are used for guidance of the catheter by combinations of push–pull and torque motions. As shown in Figure 11.1, the guide wire is inserted into the selected branch of a blood vessel and approaches the target site. A catheter is then introduced over the guide wire and the guide wire is pulled out leaving the catheter inside the vessel [2]. Guide wires of different tip shape, flexibility and size have been developed for different navigation needs and applications. Larger endoscopes equipped with imaging sensors at the tip can be manipulated with visual feedback. For smaller catheters, doctors navigate by observing the position of the catheter tip using X-ray fluoroscopy and angiography [2]. However, the acquired images are twodimensional (2-D) and lack detailed information about the vessel wall. In addition, the long distance (~1 m) between the operator’s hand and the tip of the catheter/guide wire demands a high level of skill from doctors. Potential limitations of the current catheter and guide wire designs include long procedural time, lumen or vessel wall damage and the subsequent medical complications. These issues become more critical when dealing with narrow and complex passages, such as blood vessels of the brain and tertiary bronchi of the lung. Recent advanced catheter designs exploit active tip bending for more controllable and efficient minimally invasive medical procedures. Important considerations in developing and implementing active bending catheters include the needs for miniaturization, reasonable fabrication costs, safety and the appropriate mechanical properties. These parameters are discussed in more detail in the following section.

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Blood vessel Guide wire

Balloon catheter

(1)

(3)

Inflated balloon Stenosed lesion

Treated

(4) stenosed lesion

(2) (a)

(b)

Figure 11.1 (a) Operation of conventional catheter and guide wire. (b) Process of catheter intervention in a blood vessel for treatment of a stenosed (partially occluded) lesion using a balloon catheter [2] (Reprinted with permission from Proceedings of the IEEE, Biomedical microsystems for minimally invasive diagnosis and treatment by Y. Haga and M. Esashi, 92, 1. Copyright (2004) IEEE).

11.3 Catheter Design Challenges A successful catheter design comprises a biocompatible, microsized, low cost device with a certain optimum structural rigidity, which can be actuated under safe operating conditions. 11.3.1

Biocompatibility

In designing active catheters, biocompatibility is the first requirement. Active catheters are required to be made of materials that do not have toxic or injurious effects on biological systems. There are, however, cases of the active element containing nonbiocompatible features, in which encapsulation with a biomaterial is recommended. Formation of a biolayer on the surface of the device also needs to be investigated, since it may affect the performance of the active element. In addition, the operational conditions, such as temperature, voltage, current and power, should be carefully considered in the design in order to prevent any harm to the patients and the users. According to a study on safe current limits, a current flowing from the chassis of the medical device to the ground (leakage current) of less than 50 mA results in a very low probability of injury to the patient [36]. The allowable leakage currents in medical devices made in the United States are generally set by the American National Standard Institute (ANSI). The heart is particularly sensitive to currents, and the American Heart Association has guidelines to be considered in the design of active catheters [46]. The acceptable limits are still being debated and regulations change over time.

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Biomedical Applications of Electroactive Polymer Actuators

The United States Food and Drug Administration (FDA) regulations also require appropriate precautions to be taken before releasing any medical product to the market [33]. It has established classifications for approximately 1700 different generic types of devices assigned to regulatory classes I, II and III based on the level of control necessary to assure the safety and effectiveness of the device [33]. This classification is risk based, where Class I includes devices with the lowest risk and Class III includes those with the greatest risk. FDA classified steerable catheters as Class II, which are subject to special controls. Special controls may include special labelling requirements, mandatory performance standards and post-market surveillance [33]. Further information about FDA regulations can be found at the FDA website [33]. 11.3.2

Small Size

Catheter producers are reducing the sizes of their products for two main reasons; firstly to minimize the size of the patient’s entry wound from a minimally invasive procedure, and secondly to reach the microsized arteries, which were not accessible in conventional catheterizations. 11.3.3

Low Cost

Surgical interventions where sizeable incisions are made have an associated risk of infection and also can be very expensive [33]. Minimally invasive intervention using catheterization offers patients both clinical and cost benefits [34]. Thus, it is important to design low cost active catheters to keep the cost of these procedures as low as possible. One of the challenges in producing microsized devices is their fabrication costs. As devices become smaller their fabrication procedures becomes more complicated and expensive. As an active catheter designer, it is necessary to consider choosing a cost effective fabrication method as well as low cost active element materials. 11.3.4

Structural Rigidity

A certain structural rigidity is required to satisfy a catheter’s self-sustaining specifications, depending on the application. A rigid structure enables catheter penetration along the vessels without collapsing. On the other hand, a flexible catheter tip provides better manoeuvring through twisted and convoluted paths. Thus, catheter designers often implement variable rigidity along the catheter length for optimum performance. To evaluate possible techniques for producing variable catheter rigidities, parameters affecting the rigidities need to be determined. These parameters are defined by axial, bending and torsional stiffness of the catheter (Figure 11.2), and if controlled properly will result in effective steering of the catheter. The axial rigidity is a measure of axial resistance of the structure along its length. It is found by the product of elastic modulus, E, and cross-section area, A (i.e. EA). Torsional rigidity relates to resistance to twisting and is found by the product of shear modulus, G, and polar moment of inertia, J (i.e. GJ). Flexural rigidity is a measure of the resistance of the bending deformation of a beam and is the product of elastic modulus, E, and area moment of inertia, I (i.e. IE).

Steerable Catheters Axial movement

233

Torsion

bending

Figure 11.2 Catheter function (adapted from [35]).

The axial (EA), torsional (GJ) and flexural (EI) rigidities for tubular structure of uniform composition are: EA ¼ EpðrO2

ri2 Þ;

GJ ¼ G

pðrO4

ri4 Þ 2

;

and EI ¼ E

pðrO4

ri4 Þ 4

where ro and ri are the outer and inner radii, respectively. As mentioned earlier, catheters are used in a variety of applications and the required mechanical rigidity depends on the particular applications. For instance, cardiovascular guide wire/catheter structures should have sufficient axial rigidity along the length of the blood vessels for penetration purposes and high torsional rigidity to provide a controlled twist for proper manipulation [31]. The end of the structure, however, may possess low flexural rigidity in order to move in the tortuous sections of the cardiovascular system. On the other hand, mapping cardiovascular catheters, which are designed for stimulating and monitoring electrical activity, are rigidly constructed [32]. A mapping catheter’s rigidity makes it better suited to the turbulent conditions of the interior of the heart than the flexible, guide wire directed catheters used for other interventional procedures [32]. As another example, the tips of catheters used for contrast agent delivery are specially shaped to allow the interventional radiologist to thread the catheter into side branches of arteries. The tips can be either semi-rigid or highly compliant, depending on the end use [32]. Thus there are large ranges in rigidities for existing cardiovascular catheters and the flexural and torsional rigidities differ considerably for catheters designed for different purposes [31]. The flexural rigidity of cardiovascular catheters varies from 22  10–6 to 747  10–6 Nm2, the torsional rigidity varies from ~22  10–6 to 1400  10–6 Nm2, and the axial rigidity was measured to be ~450 N [31]. In addition, guide wires have been shown to add flexural rigidity even to already rigid catheters [31]. As catheters become more advanced, the necessity to modify these rigidities becomes more evident. It is also very important to consider the varied rigidities along the catheter structure when designing an active bending catheter. An active element embedded on passive catheters (such as a conjugated polymer for controlled bending) needs to satisfy the catheter rigidity requirements for the desired application. To design a passive catheter with different rigidities along the catheter length two methods are used; varying wall thickness and incorporating reinforcing fibres in the wall using a method called two-dimensional tubular braiding [31]. 2-D tubular braiding is a fibre deposition process in which a number of fibre strands are intertwined around a tubular

234

Biomedical Applications of Electroactive Polymer Actuators

mandrel producing a stiff composite [31]. In the case of the braiding method, changes in the catheter rigidities are achieved by varying the braid angle of fibre reinforcement along the length during the braiding process. Thus, the catheter geometry does not need to be changed, resulting in a more versatile method. To convert passive catheters to actives ones, the designer should carefully account for the variable rigidities so that incorporation of the biocompatible active elements does not fail the rigidity requirements of the catheters. Alternatively, by embedding distributed active elements all along the catheter length, and controlling the bending of each segment individually, may waive the catheter rigidity requirements because the elements can, in principle, be used to control insertion without the need for applied forces and torques from the outside. However, a sufficient axial rigidity still is needed to prevent the catheter from buckling during penetration through the vessels. On the other hand, the active element may not be able to produce large bending angles during passage through tortuous paths, and thus a low flexural rigidity may still be required at the tip of the catheter for some applications. For catheters meant to be operated in close proximity to the heart, high flexural, torsional and axial rigidities are required in order to withstand the forces exerted by the rapid blood flow. Therefore, the embedded active element needs to be designed so that the resulting active catheter maintains high rigidity.

11.4 Active Steerable Catheters Current steerable catheters used in clinical practice employ a push–pull control mechanism at the handle, where a wire is configured to be pushed or pulled along a longitudinal axis to bend the catheter tip. The principal drawbacks of these steerable catheters are that their steering ability is dependent on the probe’s overall morphology and their diameter is typically significantly large. Recently, more advanced steering mechanisms have been developed. In this section steerable catheters using different types of actuators are first briefly reviewed before conjugated polymer based actuators are discussed in more detail. 11.4.1

Non-EAP Based Steerable Catheters

Micromotors installed at the tips of catheters had been used in ultrasonography [3, 4]. However, their size is relatively large (diameter of 1.9 mm [4]) and their fabrication process is expensive and complex. Shape memory alloy (SMA) coil actuators have also been developed for steerable endoscopes [5–8]. The catheter is controlled from outside the body and moves like a snake using distributed SMA coil actuators [35]. The SMA coil contracts when heated above a certain transition temperature by the application of an electrical current. SMA actuators are potentially able to provide a large degree of bending but there are some problems, such as response delay, leakage of electrical current and operating temperature [9]. Hydraulic mechanisms have also been used to control the bending motion of the catheter, where positioning of the catheter in a blood vessel is realized by varying the size of inflatable balloons installed at the tip of the catheter [2]. The balloons are inflated individually using electro-thermally controlled microvalves and the degree of catheter bending is controlled by the size of the balloons [2]. This mechanism is cumbersome and controlling the microvalves is slow; hence it is not suitable for many applications.

Steerable Catheters

11.4.2

235

EAP Based Steerable Catheters

Electroactive polymers (EAP) are a new generation of actuators which provides new approaches to propulsion and manoeuvrability. Polymers such as Conjugated (or conducting) Polymers (CP), Ionic Polymer–Metal Composites (IPMC) and Shape Memory Polymers (SMP) offer low actuation voltages, moderate to high strains and ease of preparation, making them appropriate materials for application in active catheters. Each of the sections that follow briefly describes the properties of one of the EAP actuator technologies, and then presents the properties of the catheters and related medical devices demonstrated to date, along with the potential for, and challenges involved in, further development. 11.4.2.1

Ionic Polymer–Metal Composite (IPMC) Based Steerable Catheters

The Ionic Polymer–Metal Composite (IPMC) is a type of EAP actuator whose use in active catheters has been studied the most [19–22, 27, 28]. An IPMC consists of a polymer electrolyte sandwiched between two thin metal layers and typically has an elastic modulus of 0.1 GPa [10]. Common polymer electrolytes are perfluorinated alkenes with anionicgroup-terminated side chains, such as Nafion and Flemion or styrene/divinylbenzene based polymers with ionic groups substituted from phenyl rings [10]. Deflection of the electrode/polymer/electrode structure towards one of the metal electrodes occurs as a result of applying potential across the two electrodes, where cations travel to the cathode side of the IPMC along with water molecules, which along with reorganization of hydrophilic clusters leads to polymer expansion. The strain in the cathode layer induces stresses in the rest of the polymer matrix, resulting in a fast bending motion towards the anode. After the immediate bending response, there is a slow relaxation towards the cathode. Reversing the applied potential inverts the bending. The application of an alternating voltage leads to bending vibration at the same frequency as the applied voltage [10]. IPMC actuators exhibit a typical strain of 0.5 %, strain rate of 3 %/s and a typical stress of 3 MPa. They are actuated at potentials of <10 V [42]. The performance of these actuators has been improved by using various combinations of cations [23–25] and different types of electrodes, such as platinum–copper [26]. In this chapter a brief overview of different designs and test procedures using this type of actuators is presented, as similar approaches could also be employed in conjugated polymer driven devices. IPMC based steerable catheters are further described in another chapter of this book. S. Guo et al. have proposed a microactive guide wire incorporated into the lumen of a catheter using an IPMC actuator [19]. The IPMC actuator is made from a film of perfluorosulfonic acid polymer chemically plated on both sides with platinum [20]. The active catheter was operated in saline solution over a range of temperatures and within tubes with a similar diameter and flow rates to blood vessels (2 mm inner diameter, turning angle of 45–95° and at flow rates of 50 ml/min and 650 ml/min [20]). By controlling the applied voltage (<1 V) the IPMC actuator was bent into a circular arc. The operating time with the active guide wire was half that needed to guide a passive catheter, indicating that it can improve the effectiveness of traditional procedures for intracavity operations [20]. K.Onishi et al. proposed a tube of IPMC as a microactuator for active catheters [21]. Their suggested device is a perfluorocarboxylic acid (ion exchange polymer) tube

236

Biomedical Applications of Electroactive Polymer Actuators

which is electroplated with gold on both the external and internal sides. The external gold layer is divided into four electrodes by longitudinal laser cutting. By choosing four electrodes and varying the applied voltage, the tip of the catheter can be bent in any direction [21, 2]. W. Jong et al. modelled, fabricated and tested an IPMC strip actuator for inclusion into a low cost active tip bending system for a catheterscope [22]. Platinized Nafion 117 membranes with thicknesses of 170 mm were employed. The authors studied the force generated by the IPMC actuators and found that additional force may be required in order to overcome bending resistance from the catheterscope sheathing, capillary forces of attraction within wet ducts and viscous fluids. Encapsulation of the design was suggested to avoid gas production from hydrolysis [22]. Y. Wang et al. have developed a hand-held IPMC driven scanning optical coherence tomography (OCT) probe for oral and skin imaging [28]. An optical fibre was directly attached to an IPMC actuator such that the bending vibration of the IPMC generates a scanning movement of the optical beam. The optical fibre is scanned by 3 mm at 1 Hz using an applied voltage of 2 V [28]. This design is suitable for oral and skin imaging applications; however, a more sophisticated application of OCT probes in in vivo imaging requires fibre optic scanning rates of 10–30 Hz in order to generate real time images. Real time imaging is required when operating inside a nonstationary environment, such as the coronary arteries. During some surgical procedure such as radio frequency (RF) ablation of tissues inside the heart or puncturing through blood vessel blockages, real time imaging may help guiding the surgeons to perform the operations with better accuracy. Fast scanning catheters for in vivo OCT imaging is an attractive application for EAPs, and early work in this area using conjugated polymer actuators is described below. 11.4.2.2

Shaped Memory Polymer (SMP) Based Steerable Catheters

Shaped memory polymers (SMPs) are another type of actuator used in active catheter application. SMPs can be formed into a primary shape then deformed into a stable secondary shape, and on controlled heating will return to their primary shape. Shaped memory polymers are of great interest in biomedical applications. Polyurethane, in particular, has been used in applications such as implantable devices and catheters mostly due its biocompatibility [29]. A catheter based SMP microactuator coil using polyurethane has been developed by Melodie F et al. to capture and remove the thrombus (blood clot) and restore blood flow [29, 30] as an alternative to the conventional pharmacological treatment of ischemic stroke. The SMP actuator mounted on an optical fibre is delivered in its secondary straight rod shape to the vascular occlusion through a catheter distal. The SMP actuator is then laser heated and transformed to its primary corkscrew shape to capture the thrombus. The device is finally retracted and the blood flow is restored. The length of the SMP actuator in its straight form is 4 cm and its diameter is 350 mm. Full transformation from straight rod to corkscrew occurs within 5 s, and the required laser power is approximately 1000 mW. For operation in blood, however, higher laser power is required due to the cooling effect of the blood flow and smaller corkscrew diameter (<3 mm diameter) is desired to operate in more deeply-seated vessels, where occlusions are most likely to occur [29].

Steerable Catheters

11.4.3

237

Conjugated Polymer Based Steerable Catheters

Conjugated polymer actuators are relatively new types of actuators, whose application to active catheters and cochlear implants is being actively investigated. Conjugated polymers exhibit electrochemically controllable electronic conductivities, which often reach 104–105 S/m [41]. The most common conjugated polymers used as actuators are polypyrrole (PPy) and polyaniline (PANi). The elastic moduli of these polymers are typically >0.2 GPa and can reach 7 GPa in reinforced fibres [42]. These polymers are typically disordered semiconductors in their neutral state. The effective band gap is reduced in a process known as doping [11]. Doping involves the addition (oxidation) or removal (reduction) of charges from the polymer chain resulting in creation of states in the band gap and leading to nearly metallic conduction. H N

H N

H N

v

N H

N H

N H

x

Contracted (reduced state)

Expansion

–2e– +2A–

+2e– –2A–

Expanded (oxidized state) H N

Discharging Charging Contraction

Removal of ions Insertion of ions

H N

H N N H

N H

–

N H

–

A

x

PPy

Catheter

A (a)

PPy

(b)

Figure 11.3 (a) Electrochemical redox cycle for polypyrrole with mobile anions. A– represent anions, e– electrons [14]. (b) Stress distribution upon insertion and removal of ions from the polypyrrole layer. If small cations are present with large anions the direction of ion flux and actuation is inverted (adapted from [49]).

Such a change in oxidation state in polypyrrole, induced electrochemically inside an ionic electrolyte, is illustrated in Figure 11.3a. In the process of changing oxidation state, dimensional changes are observed. The mechanism of the dimensional change is illustrated in Figure 11.3b, where a passive material (i.e. catheter) is coated with conjugated polymer electrodes on both sides forming a trilayer structure. The trilayer is seated in an electrolyte solution containing mobile negative ions and large immobile positive ions. An alternating voltage is applied across the two polymer electrodes resulting in alternating oxidation and reduction of the polymer electrodes. As shown in Figure 11.3b, during oxidation of the right hand side polymer electrode mobile negative ions enter it from a surrounding electrolyte to balance charge. The negative ion insertion results in expansion of the right hand side polymer structure. Simultaneously, the left hand side polymer electrode is in the reduced state, where mobile negative ions exit the polymer to the surrounding electrolyte, which results in contraction of the left hand side polymer structure. Expansion on one side and contraction on the other side induces a stress gradient on the polymer/catheter interfaces and causes the whole structure to bend in one direction. This process reverses itself in the second half cycle of the alternating voltage, resulting in catheter bending in the other direction.

238

Biomedical Applications of Electroactive Polymer Actuators

The electrolyte can be a liquid, a gel or a solid. Solid polymer electrolyte is a complex of high molecular weight polymers and metal salts and a gel is a liquid solution of metal salts trapped in a cross-linked soluble polymer matrix [12]. Conjugated polymer actuators are good candidates for driving active catheters, due to their low voltage (~2 V), high strain (typically 2 % or greater), low cost and biocompatibility [42]. The relatively high stiffness of these materials (elastic moduli of ~1 GPa [42]), also makes them suitable for use as active catheter drive elements. 11.4.3.1

Intravascular Microcatheters Steered by Conjugated Polymer Actuators

Della Santa et al. have proposed two different design concepts using conjugated polymer active electrodes and a solid polymer electrolyte (SPE) in order to transform passive catheters into active and steerable ones [13]. These authors suggest fabricating a 30 mm long catheter tip with external and internal diameters of 0.8 mm and 0.4 mm respectively. This active catheter is designed as an intravascular catheter. Such devices are typically expected to produce a bending angle of at least 20°, have a bending stiffness of at least 5 N/m and a response time of the order of seconds. A device with these characteristics is able to negotiate through arteries with path angle approaching 20°, and to maintain its shape in the presence of pressurized blood inside [13]. The first design is composed of conjugated polymer actuators in the form of thin strips, which are inserted inside the catheter walls (Figure 11.4a). Each actuator is built up as a CP/SPE/CP sandwich structure; the expansion or contraction of the CP layer makes the multilayer bend, and this bending is transferred to the structure of the catheter [13]. In the second configuration, the wall of the catheter is fabricated from an active composite of CP fibres embedded in an SPE elastometric matrix (Figure 11.4b). Figure 11.4c illustrates the operation of the second proposed catheter, in which two opposite electrochemical stimuli are applied: one causes contraction of the fibres in the upper half of the structure, while the other causes expansion of the lower fibres, leading the catheter tip to bend upwards in this case [39]. Sheath

Sheath

CP

Composite

SPE (a)

(b)

(c)

Figure 11.4 Schematic of the device: (a) first design consists of CP/SPE/CP sandwich structure inserted inside the catheter wall [39]; (b) distributed actuator fibre þ SPE; (c) functioning device (adapted from [12]).

As mentioned in Section 11.2.4, the active catheter needs to be designed so that the rigidity requirement for the desired application is fully satisfied. The axial rigidity of the structure needs to be high enough to avoid the catheter collapsing during the insertion. The flexural rigidity, which relates to bending stiffness also, needs to be low enough to allow the required degree of bending (i.e. 20° was set in this design).

Steerable Catheters

239

In order to study the feasibility of the proposed designs shown in Figure 11.4, a finite element simulation was performed by Mazzoldi and colleagues [13]. To simulate the structure’s mechanical behaviour, parameters which quantify the expansion or contraction of the CP due to electrochemical excitation need to be determined. These parameters include Young’s modulus of the passive catheter sheaths, the conjugated polymer and of the SPE, and the CP electrochemical strain. In the simulation reported [13], a commercial biocompatible polyethylene tube with an elastic modulus of 140 MPa was assumed to form the internal and external sheaths and to act as the electrochemical insulator. The two active configurations shown in Figure 11.4 were modelled. In the first configuration, the CP/SPE/CP sandwich structure is inserted inside the catheter walls. Polypyrrole doped with benzensulfonate anions (PPy/BS–) with an elastic modulus of 450 MPa and active strain of 1 % was assumed to be the CP element, and a layer containing Cu(ClO4)2 with a modulus of 45 MPa was considered as the SPE. In the second configuration, the walls of the catheter are assumed to be made of the CP fibres/SPE matrix composite material. The CP fibres may be PPy or PANi extruded microfibres. In the simulation the fibres were assumed to have the same active and passive properties as their film form. The overall mechanical characteristics of the composite structure are as follows: Young modulus 247 MPa, shear modulus 37 MPa and electrochemical strain 1 % [13]. According to the simulation results, the second configuration (Figure 11.4b) offers a better performance both in terms of bending degree (17.2°) and bending stiffness (0.81 N/m) than the first configuration (11.7° and 0.19 N/m) [39]. The performance of the first configuration in terms of rigidity can be improved by increasing the stiffness of the SPE; Mazzoldi et al. suggested charging the material with inert powders or through irradiation in order to increase the SPE stiffness. However, the second configuration seems to enable better rigidity without affecting the amount of catheter bending. This increase in rigidity can be obtained by increasing the fraction of composite material inside the walls of the catheter [39]. In order to further improve the rigidity, Mazzoldi proposed incorporating polyaniline fibres doped with perchlorate ions (PANi/ClO4) as the conjugated polymer element, which results in an elastic modulus of 1.5–4.5 GPa [12]. The thickness of the CP is an important factor determining the actuator response time, and since the fibres can be much smaller than the films, they may also be faster [12]. CP fibres with diameters of less than 10 mm are likely needed in order to obtain the required actuation speed [12]. In the optimum configuration (Ef ¼ 4.5 GPa, Em ¼ 45 MPa), simulation results gave a degree of bending of 28° and a bending stiffness of 5 N/m [12]. Only in this case can the catheter lumen host the fibre optic bundles (estimated elastic modulus: 10 GPa) without considerably losing its bending capability. The simulation results for an optical fibre embedded active catheter designed in [12] indicate a bending angle of 24° and a bending stiffness of 15 N/m.The active catheter with these characteristics is able to reach through arteries with a path angle of less than 20°; however, a more complicated design capable of producing a larger angle of bending is needed in order to access tortuous regions of cardiovascular systems. 11.4.3.2

Conjugated Polymers in OCT Fast Scanning Catheter

The authors of this chapter are investigating active catheter designs similar to those of Mazzoldi, De Rossi and colleagues that are described in the previous section, but employed to rapidly scan the tip of the catheter to enable high-resolution 3-D imaging via optical

240

Biomedical Applications of Electroactive Polymer Actuators

coherence tomography. In vivo imaging of internal body tissues using OCT catheters has a number of potential applications, including the diagnosis and monitoring of the progression of vascular diseases [47]. Forward-looking images from the inside of the blood vessels can be useful for navigating catheters during interventational procedures. Due to the size and the torturous nature of the blood vessels, fibre bundle based systems are typically not compatible for clinical applications in the brain or heart. Regardless of the operating wavelength (visible or near infrared), red blood cells produce significant multiple scattering that flushing with saline is required to obtain clear images [2]. Optical Coherence Tomography is an emerging imaging technique that permits three-dimensional highresolution visualization of subsurface tissue, although saline flush is still required when performing intravascular imaging [15]. It can serve as an important diagnostic adjunct, enabling the detection and the monitoring of changes in tissues, may be useful for assessment of vascular disease and cancer tissue progression [16] and surgical guidance. OCT performs depth resolved imaging by sending wide band near infrared light into tissue and observing the backscattered light interferometrically [16]. One of the common designs currently employed by intravascular OCT uses a rotating fibre driven by an external motor. The fibre tip contains miniature focusing and beam bending optics such that rotation of the tip provides circumferential scanning of the optical beam in a side viewing geometry [17]. Physical displacement of the catheter through the arteries, coupled with the rotational motion, provides 3-D cylindrical scanning. A limitation of this approach is the lack of imaging information distal to the tip of the catheter (i.e. in front of it), which may be addressed by a forward-viewing design. Forward viewing will enable imaging of occlusions for example. The authors of this article have designed a polypyrrole actuator for scanning an optical fibre in two dimensions. The desired scan area is 1 mm2, at a relatively high speed (10–30 Hz per line scan) and over ~50 000 cycles for intravascular OCT imaging application. The design is composed of a four-electrode polymer actuator, which is fabricated directly on a commercial catheter tip. The catheter used to demonstrate the bending is Micro Therapeutics Inc. (Irvine, CA) UltraflowTM HPC (0.5 mm OD, 0.28 mm ID) where the final 25 mm is coated with active polymer. The complete device is a catheter whose active tip is encapsulated inside a structure containing an ionic electrolyte (Figure 11.5). The device is electrochemically actuated inside the ionic electrolyte in two dimensions

Encapsulation

Electrolyte Active catheter

Lens

Optical fibre

Imaging plane

Figure 11.5 Schematics of the active OCT imaging catheter, where an optical fibre is scanned by the polymer actuator and a lens focuses the light at the imaging plane. The reflected light is collected by the same optical fibre and processed by the OCT system.

Steerable Catheters

241

using voltage changes of <1 V per polymer electrode. Actuation is associated with the insertion or removal of ions to or from the polymer as was explained earlier. The encapsulation ensures that currents needed to drive the actuation are isolated from nearby tissue, and provides an optimal electrolyte for achieving large and fast actuation. A lens is embedded at the end of the encapsulation tube, which can amplify the angular scanning range. The active element fabrication involves chemical deposition of polymer onto the catheter followed by an electrochemical deposition. Chemical deposition is performed by polymerization of pyrrole monomer in the presence of an oxidizing agent. The oxidizing solution contains 0.413 ml of 12.1 M HCl and 1.2 g of FeCl3 in 50 ml of water. The monomer solution contains 0.1 M of pyrrole. Deposition is achieved by alternately dipping the sample in each of the two solutions until the desired thickness of polymer forms [16]. This process results in a thin polymer with a conductivity of ~103 S/m, which is sufficient to be used as a seed layer in electrochemical deposition. Polypyrrole films are grown from a solution of 0.06 M pyrrole monomer and 0.05 M tetraethylammonium hexafluorophosphate and 1 vol-% distilled water in propylene carbonate, following the procedure of Yamaura [14]. Polypyrrole is deposited galvanostatically on to electrically conductive substrates at current density of 1.25 A/m2 and at temperatures between –30 °C and –45 °C. Four longitudinal stripes are then cut into the polymer coating via laser ablation to produce four electrodes (Figure 11.6).

Plastic Substrate

Spiral Skeleton

1 mm (a)

4 AMAs along the Tip

(c)

Differential Actuation of AMAs for Bending

Outer Diameter < 1 mm

(b)

(d) 600 µm

Electron Image 1

Figure 11.6 Illustration of the design and prototypes of four-segment conjugated polymer electrodes for two-dimensional motion.

242

Biomedical Applications of Electroactive Polymer Actuators

Effective patterning of the polypyrrole electrodes, shown in Figure 11.6, is critical to the success of this design. Patterning of polypyrrole electrodes has previously been achieved using photolithography [43], but this method is not practical for nonplanar geometries. Instead, laser micromachining was employed. It is a mature technology for shaping tubular structures such as laser cut stents. Excimer laser micromachining, in particular, is very well established as a tool for machining of polymers. Excimer laser ablation at 248 nm was used to pattern polypyrrole on the coated catheters (Kr:F excimer laser, GSI Lumonics plus beam expansion optics, a photomask, focusing optics and an xyz programmable stage). An important benefit of using ablation is that there is no thermal damage and minimal debris, as is evident in the clean cut depicted in Figure 11.7.

Figure 11.7

A 50 m channel in the polypyrole created by excimer laser ablation.

A preliminary image was obtained by inserting a ball-lensed fiber optic into the polypyrrole based active catheter. The fibre was connected to a commercially available OCT system (Figure 11.8a). The scanning angle using the ball lens is smaller than using a larger lens, such as gradient index lenses [48]. The primary demonstration of the device was performed with a speed of 10 Hz. In order to achieve this frequency response a relatively large overvoltage was needed, with the total applied voltage reaching 10 V. The acquired image was comparable to the images taken using the conventional motor based fibre scanning method. However, the number of scanning cycles was limited due to short polymer lifetime at high actuation voltages. Therefore, design modification is required in order to achieve both high speed of actuation and the desired polymer life cycle of >50 000. A dynamic electromechanical model was developed to simulate the actuation behaviour of the active catheter, and was used to study the feasibility of achieving high speed actuation (10–30 Hz) appropriate for OCT imaging. The model consists of a flexible tube (catheter) with inner and outer diameters of b ¼ 280 mm and a ¼ 500 mm with 10 mm thick polypyrrole coating as the active element. According to the model, a polymer thickness of ~10 mm results in maximum tip displacement at high frequencies (10–30 Hz). The Young’s modulus

Steerable Catheters

243

To OCT system

electrode

electrode

NaPF6 ~1mm

Optical fiber+lens

Sample

(a)

(b)

Figure 11.8 (a) Schematic of the actuator set up for OCT imaging. (b) Preliminary image of an IR card obtained by scanned OCT fibre.

of a 10 mm thick electrochemically grown polypyrrole film and a commercially available catheter were measured to be 0.3 GPa and 75 MPa respectively. The polypyrrole active strain, ", was measured to be ~2 % during actuation of films between –0.5 and þ0.5 V vs Ag/AgCl reference electrodes in aqueous 1 M NaPF6 electrolyte. This voltage range is well within the safe actuation voltage range [40]. The actuator length is limited to a maximum of between 10 and 35 mm (set for the specific intravascular application). As mentioned earlier, the active catheter is encapsulated inside a rigid encapsulation, thus the rigidity requirement of the catheter is satisfied as long as the encapsulation possesses rigidities of common cardiovascular catheters (i.e. flexural rigidity of >22  10–6 Nm2, torsional rigidity of >22  10–6 Nm2 and axial rigidity >450 N [31]). The rate of actuation of polypyrrole is proportional to the rate of ion insertion, and hence to the current. The current can be limited by the both ionic and electronic resistivities, diffusion coefficients and the capacitance of the electrodes. For a long device the total polymer resistance along the length of the catheter, Rppy, can significantly limit the 2 rate [41], with the predicted time constant being t ¼ Rppy C ¼ Cvel [41], where se is the electronic conductivity of the polymer and Cv is the polymer volumetric capacitance. For a thick device the ionic resistance through the thickness, tp, is the rate limiting factor, with t2

the time constant of t ¼ Dp [41], where D is the effective diffusion coefficient. In electrochemically deposited PPy grown using Yamaura’s method, a diffusion coefficient of D ¼ 5.5  10–11 m2/s, and electronic conductivity of selectronic ¼ 1  104 S/m were found. These values are fed into models to help calculate the ion insertion rate and hence predict the actuation speed. According to the model, the desired high speed scanning (30 Hz) cannot be achieved using these values for the designed geometry. Thus, a geometry optimization was performed using a dynamic electromechanical model, in order to

244

Biomedical Applications of Electroactive Polymer Actuators

improve the performance of the device at high frequency operation. The model suggests mechanical amplification and electromechanical parameter improvement to achieve the required 1 mm displacement at high frequency (Table 11.1). Table 11.1 "¼4%

Suggested parameters for high speed actuation

D ¼ 9  10

Design Active Length (mm) Displacement (mm)

11

m2/s

selectronic ¼ 5  104 S/m @ 30 Hz 35 1.0

@ 10 Hz 13 1.0

One approach to achieving the needed displacements at high scan rates is to leave the tip of the catheter uncoated, and simply coat the catheter further down, away from the tip. Thus, mechanical amplification is provided by adding a passive length to the active element. Another approach is to increase speed by increasing diffusion coefficient, or reducing the thickness of the active layer [44, 45]. The fabrication process can also be modified in order to obtain highly porous polymer, which helps increase the ion insertion rate. This may, however, reduce the structural rigidity, and therefore careful considerations are required to maintain the specified rigidity as well as high ion diffusion rate. Electronic conductivity of the active element can be enhanced by coating the polymer with a thin metal layer. The metal layer adds to the stiffness of the device and may decrease the degree of catheter bending. The primary challenge to achieving an effective polypyrrole driven catheter is to demonstrate high speed with reasonable liftetime. Geometry optimization is needed to successfully complete the project. 11.4.3.3

Conjugated Polymers in Cochlear Implants

The use of conjugated polymer actuators in cochlear implants is being studied by the Intelligent Polymer Research Institute [18]. The cochlear implant device is a multi-channel array of electrodes that is brought into contact with the hearing nerves and electrically stimulates these nerves to restore hearing in deaf people. Insertion of an array of electrodes into the cochlea presents a challenge that is similar to the guidance of a catheter through the body vessels. Figure 11.9 shows one of the most common cochlear implant electrode arrays (i.e. Nucleus 24 contour). The electrode is held straight by a thin platinum stylet in a lumen in the electrode prior and during insertion. The stylet is partially removed during the insertion and fully removed after the full insertion [18]. This method of electrode insertion is associated with a high risk of damaging auditory neurons [18]. Therefore, a steerable actuator for guiding electrode insertion is needed. It is also desirable to position the electrode array close to the ganglion cells in the modulus for effective hearing restoration [37]. Zhou and his colleagues at the University of Wollongong developed a conjugated polymer based actuator for accurate and controllable insertion of the device into the inner ear. This technology allows optimal hearing restoration by providing close proximity of the array to the auditory neurons and less auditory damage. Zhou fabricated a prototype actuator and attached it to the back of a Nucleus 24 Contour

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Figure 11.9 (a) Nucleus 24 contour cochlear implant electrode array (Reprinted with permission from Synthetic Metals, Actuators for the cochlear implant by Zhou, D., Wallace, G.G., Spinks, G.M. et al., 135–136, 39–40. Copyright (2003) Elsevier). (b) A diagram showing a standard nucleus electrode array inserted around the spiral of the first turn of the cochlea to stimulate the auditory nerve fibres [37] (Reprinted with permission from Medical Engineering & Physics, Evaluation of trajectories and contact pressures for the straight nucleus cochlear implant electrode array – a two-dimensional application of finite element analysis by Chen, B.K., Clark, G.M. and Jones, R., 25, 2, 141–7. Copyright (2003) Elsevier).

0

PPy PPy Polymer in ionic liquid electrolyte within the pores of PVDF

0

Figure 11.10 Diagram depicting the cross-section of double polypyrrole actuator on PVDF that is used to deflect the cochlear implant electrode array shown in Figure 11.9.

electrode (Figure 11.10). The actuator consists of a two 30 mm thick polypyrrole layers electrochemically deposited on a 100 mm thick, platinized porous polyvinylidene fluoride (PVDF) membrane. The PVDF membrane contains propylene carbonate solvent with 0.25 M tetrabutylammonium hexafluorophosphate [18]. The electrode can be steered and inserted in a controllable manner by applying a voltage across the polymer electrodes. Electrode arrays for cochlear implants are usually inserted 22–30 mm from the opening to the inner ear into the cochlea, with a total deflection of 380–420° [38]. The actuator [18] was able to achieve a bending angle of more than 180° on the PVDF membrane itself from a 2.4 % strain. However, the bending angle is less than 180° when the actuator is attached to the implant electrode due to high stiffness of the electrode [18]. This is not sufficient to steer the electrode through the spiral structure of the cochlea to achieve the required insertion depth of 22 mm. Design modification is required to further improve the angle of bending and hence the insertion depth. One possible way to increase the bending angle is to decrease the cochlear implant electrode rigidity. The actuator’s electrochemical strain might also be increased by using an electrolyte containing larger mobile ions to generate a

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higher angle of bending. The distribution of several CP actuator segments along the length of the electrode, each capable of generating a desired degree of bending would result in a more controllable and deep insertion of the electrode.

11.5 Discussion and Conclusion Active catheterization will pave the way for minimally invasive medical diagnosis and treatments that involve less risk and pain for patients and less complication for doctors. Electroactive polymer (EAP) actuators are good candidates for use in active catheters due to their biocompatibility, low cost, large strain, low actuation voltage and ease of fabrication. However, electrochemical actuation of these devices involves using ionic electrolytes which require either encapsulation or perhaps direct use of internal fluids. EAP actuators principally used in steerable catheters are ionic polymer–metal composites (IPMCs) and conjugated (or conducting) polymers (CPs). IPMC based catheters have been widely studied. These actuators can generate large displacements at relatively low voltages (<10 V) and are typically faster than conjugated polymers; however, their manufacturing process is often relatively expensive and, unlike conjugated polymers, additional energy is usually consumed to hold the actuator in place. CPs offer higher stiffnesses than IPMCs, an attribute which is often important in catheter design. Their efficiency can be higher, as maintaining deflection does not require energy expenditure. The speed of conjugated polymer actuators is typically slow, however (only a few hertz) [42], and needs to be improved to make possible the rapid scanning used in imaging methods. Cycle and shelf life have received little attention. Cycle life is likely acceptable in disposable applications, but is likely to pose a problem in reuseable scanning devices. Implementing EAP actuators for active catheter application is still in the research and development stage. However, technological progress of these actuators along with clinical demands will advance the development of these actuators in active catheter technology.

References 1. Patel, S.R. and Caldamone, A.A. (2004) The history of urethral catheterization, Medicine and Health Rhode Island. 2. Haga, Y. and Esashi, M. (2004) Biomedical microsystems for minimally invasive diagnosis and treatment, Proceedings of the IEEE, 92, 1, 98–114. 3. Salimuzzaman, M., Matani, A., Oshiro, O., et al. (1996) Visualization of a blood vessel using a micromotor, Tech. Dig. 14th Sensor Symp. 279–80. 4. Lehr, H., Ehrfeld, W., Hagemann, B., et al. (1997) Development of micro and millimotors, Min. Invas. Ther. Allied Technol., 6, 191–4. 5. Ikuta, K., Tsukamoto, M. and Hirose, S. (1988) Shape memory alloy servo actuator system with electric resistance feedback and application for active endoscope, Proceedings of the IEEE International Conference on Robotics and Automation, Philadelphia, PA, 24–29 April 1988, 427–30. 6. Reynaerts, D., Peirs, J. and Brussel, H.V. (1996) Design of a shape memory actuated gastrointestinal intervention system, Proceedings of the 5th International Conference on New Actuators, Bremen, Germany, 26–28 Jume 1996, 409–12.

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7. McCoy, W.C. (1985) Steerable and aimable catheter, US Patent 4543 090. 8. Kaneko, S., Aramaki, S., Arai, K., et al. (1995) Multi freedom tube type manipulator with SMA plate, J, Intell. Mat. Syst. Struc., 7, 3, 331–5. 9. Fukuda, T., Guo, S., Kosuge, K., et al. (1994) Micro Active Catheter System with Multi Degrees of Freedom, Proceedings of the IEEE International Conference on Robotic and Automation, San Diego, California, 3, 2290–5. 10. Madden, J.D.W., Vandesteeg, N.A., Anquetil, P.A., et al. (2004) Artificial muscle technology: physical principles and naval prospects; IEEE J. Oceanic Eng., 29, 3, 706–28. 11. Siegmar, R. (1995) One-Dimensional Metals, Springer-Verlag, New York. 12. Mazzoldi, A., De Rossi D. (2000) Conductive polymer based structures for a steerable catheter, Proceedings of the SPIE (The International Society for Optical Engineering), 3987, 273–80. 13. Della, A. Santa, A. Mazzoldi, D. de Rossi (1996) Steerable microcatheters actuated by embedded conducting polymer structures, J. Intelligent Mat. Syst. Struc., 7, 3, 292–300. 14. Madden, J.D. (2000) Conducting polymer actuators, PhD thesis. 15. Huang, D., Swanson, E.A., Lin, C.P. (1991) New York: Optical Coherence Tomography, Science, 254, 1178. 16. Yang, V.X.D., Tang, S.J., Gordon, M.L., et al. (2005) Endoscopic Doppler optical coherence tomography: first clinical experience with a novel imaging technique, Gastrointestinal Endoscopy, 61, 879–90. 17. Yeow, J.T.W., Yang, V.X.D., Chahwan, A., et al. (2004) Micromachined 2-D scanner for 3-D optical coherence tomography, Sensors and Actuators, A: Physical, 117, 2, 331–40. 18. Zhou, D., Wallace, G.G., Spinks, G.M., et al. (2003) Actuators for the cochlear implant, Synth. Met., 135-136, 39–40. 19. Gue, S., Fukunda, T., Kosuge, K., et al. (1995) Micro Catheter System with Active Guide Wire, Proceedings of the IEEE International Conference on Robotic and Automation, Nagoya, Aichi, Japan, 1, 79–84. 20. Guo, S., Nakamura, T., Fukuda, T. (1996) Microactive guide wire using ICPF actuator-characteristic evaluation, electrical model and operability evaluation, Proceedings of the IEEE IECON 22nd International Conference on Industrial Electronics, Control, and Instrumentation, Taipei, ROC, 1312–7. 21. Onishia, K., Sewa, S., Asaka, K., et al. (2000) Bending Response of Polymer Electrolyte Actuator, Proceedings of SPIE, Smart Structures and Materials, 3987, 121–8. 22. Yoon, W.J., Reinhall, P.G., Seibel, E.J., Analysis of electro-active polymer bending: A component in a low cost ultra thin scanning endoscope, Sensors and Actuators: A Physical, 133, 2, 506–17, (2007). 23. Nemat-Nasser, S., Wu, Y. (2003) Tailoring actuation of ionic polymer-metal composites through cation combination, Proceedings of SPIE, Smart structures and materials: electroactive polymer actuators and devices (EAPAD), 5051, 245–53. 24. Onishi, K., Sewa, S., Asaka, K., et al. (2001) The effects of counter ions on characterization and performance of a solid polymer electrolyte actuator, Electrochim. Acta 46 (8), 1233–41. 25. Yamaue, T., Mukai, H., Asaka, K. and Doi, M. (2005) Electrostress diffusion coupling model for polyelectrolyte gels, Macromolecules, 38, 1349–56. 26. Uchida, M., Xu, C., Guilly, M.L. and Taya, M. (2002) Design of Nafion actuator with enhanced displacement, Proceedings of SPIE 4695, Smart structures and materials 2002: electroactive polymer actuators and devices (EAPAD), 57–66. 27. Fang, B.K., Ju, M.S., Lin, C.K. (2007) A new approach to develop ionic polymer–metal composites (IPMC) actuator Fabrication and control for active catheter systems, Sensors and Actuators: A Physical, 137, 321–9. 28. Yuli, W., Bachman, M., Guann-Pyng, L., et al. (2005) Low-voltage polymer-based scanning cantilever for in vivo optical coherence tomography, Optics Letters, 30, 1, 53–5.

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29. Metzger, M.F., Wilson, T.S., Schumann, D., et al. (2002) Mechanical properties of mechanical actuator for treating ischemic stroke, Biomedical Microdevices, 4, 2, 89–96. 30. Maitland, D.J., Wilson, T., Metzger, M. and Schumann, D.L. (2002) Laser-activated shape memory polymer microactuators for treating stroke, Proceedings of SPIE (The International Society for Optical Engineering), 4626, 394–402. 31. Carey, J., Fahim, A. and Munro, M., (2004) Design of braided composite cardiovascular catheters based on required axial, flexural, and torsional rigidities, J. Biomed. Mat. Res.: B App, Biomat, 70, 1, 73–81. 32. Kalorama Information. (1996) Market for Disposable catheters and suppliers market. Kalorama Information, New York, NY. 33. www.FDA.gov. 34. Kalorama Information. (2008) The worldwide market for catheters. Kalorama Information, New York, NY. 35. Haga, Y., Esashi, M. and Maeda, S. (2000) Bending, torsional and extending active catheter assembled using electroplating, Proceedings of the IEEE Micro Electro Mechanical Systems (MEMS), 181–6. 36. Levin, M. (1994) Safe current limits for electromedical equipment and hazards to patients, J. Am. Heart Ass., Circulation, 2160. 37. Chen, B.K., Clark, G.M., Jones, R. (2003) Evaluation of trajectories and contact pressures for the straight nucleus cochlear implant electrode array—a two-dimensional application of finite element analysis, Medical Eng. Phys, 25 141–7. 38. Dormana, M.F., Loizoub, P.C. and Rainey, D. (1997) Simulating the effect of cochlear-implant electrode insertion depth on speech understanding, J. Acoust. Soc. Am., 2993. 39. Santa, D. and De Rossi, (1996) Intravascular microcatheters steered by conducting polymer actuators, Proceedings of the Annual International Conference of the IEEE Engineering in Medicine and Biology, 5, 2203–4. 40. Shoa, T., Cole, M., Yang, V. and Madden, J. (2007) Polypyrrole actuator operationg voltage limits in aqueous NaPf6, Proceedings of SPIE, Smart Structures and Materials: Electroactive Polymer Actuators and Devices, 6524, 652421. 41. Madden, J.D.W., Peter, G.A. and Hunter, I.W. (2002) Conducting polymer actuators as engineering materials, Proceedings of SPIE, Smart Structures and Materials: Electroactive Polymer Actuators and Devices, 4695, 176–190. 42. Actuatorweb.org. 43. Jager, E.W.H., Smela, E. and Inganas, O. (2000) Microfabricating Conjugated Polymer Actuators, Science, 290, 1540–5. 44. Bay, L., West, K., Sommer-Larsen, P., et al. (2003) A conducting polymer artificial muscle with 12% linear strain, Adv. Mat., 15, 310–3. 45. Hara, S., Zama, T., Takashima, W. and Kaneto, K. (2005) Free-standing polypyrrole actuators with response rate of 10.8%s-1, Synth. Met. 149, 199–201. 46. www.americanheart.org. 47. Tearney, G.J., Brezinski, M.E., Boppart, S.A., et al. (1996) Catheter-Based Optical Imaging of a Human Coronary Artery, J. Am. Heart Ass., Circulation. 94, 3013. 48. Munce, N.R., Mariampillai, A., Standish, B., et al. (2008) Electrostatic Forward-Viewing Scanning Probe for Doppler Optical Coherence Tomography using a Dissipative Polymer Catheter, Optics Letters, 33, 7, 657–9. 49. Cohen, Y.B, Electroactive Polymer (EAP) Actuators as Artificial Muscles: Reality, Potential, and Challenges, 2nd edn, SPIE Press Monograph Vol. PM136, (2001).

12 Microfabricated Conjugated Polymer Actuators for Microvalves, Cell Biology, and Microrobotics Elisabeth Smela Department of Mechanical Engineering, University of Maryland, USA

12.1 Introduction Robots that swim in the bloodstream, scissors that take a biopsy from within a blood vessel deep inside a kidney, tapping fingers that measure the hardness of cells to distinguish cancerous from normal tissue, and smart pills that dispense drugs in concentrations that change with the body’s needs – these are among the future technologies that have been envisioned by engineers and doctors for minimally invasive diagnosing and treating of patients. However, working on the microscale like this will require microscale tools (microelectromechanical systems, or MEMS) and, furthermore, the tools must be able to operate in a hostile environment: one that is aqueous, salty, and warm. Traditional machining will not be able to produce such tiny mechanical devices, but microfabrication could, provided that it takes advantage of electroactive polymer (EAP) actuators to serve as artificial ‘muscles’ to power the movements. MEMS designs that use conventional inorganic materials are too fragile [1, 2], but polymer actuators are robust [3]. Among the most compelling EAPs for the applications envisioned above are conjugated polymers, for several reasons [4]. Firstly, conjugated polymers have good overall actuation performance in terms of force, stroke, and speed. They can exert forces that are 1000 times higher than are exerted by mammalian muscle [5], so the actuators will be strong enough to carry out the required tasks, even when miniaturized. These polymers have a DC response,

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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meaning that they can be held at a fixed position, as well as a large strain rate, so they can be rapidly re-positioned. This combination permits complex maneuvers [3, 6]. This is important for devices like microsurgical tools that may need to both cut and hold onto a sample. Secondly, conjugated polymers are operated at low voltages, so there is never a risk to the surrounding tissues. Low voltage also means that they can be directly battery-powered without transformers and controlled by standard integrated circuits (ICs). In fact, these actuators can be microfabricated directly on the surface of IC chips [7] to produce truly miniaturized systems. Thirdly, their usual operating environment is an aqueous electrolyte, so biofluids provide the water and salt needed for their actuation. These materials can be operated directly in cell growth medium, for example (although it is still unclear whether they can be successfully operated for extended periods in fluids such as blood and urine). Fourthly, the speed of the actuators actually benefits from scaling down to microstructure dimensions because thinner films can be switched faster (switching speeds are limited by ion transport rates). In addition, for bending bilayer configurations, the smaller sizes also mean a smaller moment arm. Fifthly, microfabrication procedures have already been developed, as reviewed below, and a range of devices has been demonstrated. Finally, conjugated polymers are biocompatible [8]. In this chapter, some of the exciting prototype microscale biomedical devices that have been demonstrated in the laboratory are reviewed. (Only mechanical actuation is covered; molecular-releasing films for drug delivery or neural stimulation are not included.) Before that, however, the actuation mechanism and performance are reviewed, and how such devices are fabricated is then described.

12.2 Actuator Background As reviewed in this book and elsewhere, conjugated polymers change volume due to the ingress and egress of solvated ions. The ions enter and leave the polymer from an adjacent electrolyte in response to changes in the charge induced on the polymer chains by an applied voltage, the ions serving to maintain overall charge neutrality in the material. The relative change in length of the polymer (actuation strain) therefore depends on the ions, the solvent, and the potential, as well as on the molecular structure of the polymer. Strain in conjugated polymers is typically of the order of a few % in the plane of the film, and tens of % in the out-of-plane (thickness) direction [9, 10]. Both in-plane and out-of-plane strains have been exploited in microactuators, as shown below, with the former used in bilayer bending configurations that amplify the smaller strain. The speed with which the actuators can be switched between their expanded and contracted states depends on the polymer film thickness, since actuation depends on mass transport. Like the strain, speed also depends on the electrolyte and on the polymer structure, and the structure depends not only on the type of polymer, but also on the film preparation conditions. Response times for the thin films that are used in microactuators are of the order of a second, which is sufficiently fast for most biomedical applications. The high force that conjugated polymers can exert is due to their relatively high Young’s moduli (0.05–100 GPa), and they also have high tensile strengths (1 MPa to 1 GPa) [11]. Linear actuators can achieve actuation stresses of tens of MPa [12], and bending microactuators can lift tens of thousands of times their own weight [3].

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Returning to the actuation mechanism, when an oxidizing (anodic) potential is applied to a conjugated polymer, electrons are removed from the backbone, producing positive charge carriers. These are compensated by negatively charged ‘dopant’ anions. Upon applying a reducing (cathodic) potential, the electrons are returned to the backbone, and either the dopant anions exit the film or positively charged cations enter the film. In the former case, the polymer expands in the oxidized state and, in the latter case, it expands in the reduced state, but in both cases the polymer expands upon ion ingress and shrinks upon ion egress. The number of electrons removed from the polymer (the consumed charge) thus determines the number of ions in the polymer, and thereby the actuation strain. These devices are therefore analog, in that any oxidation level between the fully oxidized and fully reduced state can be specified, and thus any intermediate position of the actuator can be held. While the basic actuation mechanism has been established and confirmed, and some conjugated polymer actuators have been characterized, there are still a number of things that are not yet understood. Basic science studies are still being performed to elucidate the complex interactions among the ions, solvents, and polymer chains that occur during switching, as well as to determine how polymerization conditions determine polymer structure, and how that structure affects actuation. In addition, further basic actuator characterization and optimization are still required, and comprehensive, predictive models of actuator behavior are still in the initial stages of development. Nevertheless, enough is known to have enabled the creation of a variety of proof-of-concept microscale devices, and to begin contemplating commercialization.

12.3 Microfabrication Microfabrication is used in the semiconductor industry to make integrated circuits and has also been adapted to the fabrication of MEMS. It employs sequential film deposition and patterning steps on a planar surface; MEMS structures are built layer-by-layer in such a way that they are assembled in place, and are partially freed from the surface so that they can move in a final ‘release’ step [13, 14]. Two different processes sequences showing the steps for fabricating a metal/polymer bilayer actuator that can rotate a plate through 180° out of plane are illustrated in Figure 12.1. (The rotation is shown in Figure 12.2.) Applications of this structure are discussed in Section 12.4.1. In the first process sequence (Figure 12.1), a sacrificial layer, such as a thin film of titanium (light gray), is deposited on the surface of an oxidized silicon wafer (gray), or other insulating surface, and patterned (a). A structural layer of gold (white) that forms half the bilayer is then deposited over the entire surface, following a thin underlying chromium film that serves as an adhesion layer to stick the gold (Au) to the oxide (b). The plate material (black), such as a layer of SU8, a negative photoresist, is deposited and patterned next (c). The gold and chromium are then patterned, typically by wet etching (d), and the conjugated polymer (dark gray) is deposited onto the gold electrode (e). Notice that the gold electrode underlies both the conjugated polymer and the plate, connecting the two. The sacrificial layer is removed by under-etching, leaving the structure attached on the left, but free to move on the right (f). Note that this leaves a small step in the structure. During

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Figure 12.1 Left: fabrication sequence based on a sacrificial layer. Right: fabrication based on differential adhesion (Reprinted with permission from [15]. Copyright (1995) IEEE).

Figure 12.2 Photographs of bilayers rotating rigid plates within an aqueous electrolyte. Left: the bilayers are in the reduced state and straight, with the plates lying flat on the substrate. Right: the bilayers are oxidized and bent into a semi-circle, and the plates are rotated 180°, gold-side up. The lighter-shaded rectangles for differential adhesion are visible underneath (Reproduced with permission from [19]. Copyright (2006) Elsevier).

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electrochemical switching the bilayer bends and straightens, rotating the plate, which remains straight (g). In an alternative process sequence, thin layers of chromium and gold are deposited and patterned to leave an opening down to the oxide (a’). A structural layer of gold is then deposited (b’). This gold does not adhere to the oxide, but it does stick to the first layer of gold. The plate layer is deposited and patterned as before (c’). The conjugated polymer is deposited over the exposed gold (d’). The polymer is patterned (e’), for example by reactive ion etching in an oxygen plasma. The exposed gold is etched, freeing the structure (f’). During the first switching cycle, the structure lifts itself off the surface. The most frequently used conjugated polymer on the microscale is polypyrrole (PPy), although polyaniline (PANi) has also been used on occasion [17, 18]. Of the polypyrroles, the one that has been most often used is PPy(DBS), PPy doped with dodecylbenzenesulfonate, which is a bulky surfactant. This material expands in the reduced state due to cation ingress and contracts in the oxidized state. There are several ways to deposit conjugated polymers. Polypyrrole is insoluble and must therefore be electrochemically deposited. Polyaniline, polythiophenes, and poly(3,4ethylene dioxythiophene) (PEDOT) are soluble, and so can be spin-coated or cast onto the substrate. There are several methods for patterning the polymers. They can be electrochemically deposited onto a patterned electrode, as shown in the first process sequence, so deposition and patterning are achieved in a single step. Alternatively, they can be deposited within a template, for example onto an electrode that is partially covered with a patterned layer of photoresist. The polymer only grows on the exposed metal, and the resist is subsequently removed. For best results, the method shown in the second process sequence is used. The polymer is deposited over the entire electrode, then covered with an etch mask layer, such as patterned photoresist, and etched. This method produces the most uniform film thickness since there are no electrode edges, and hence no thickness variations due to chemical loading effects or electric field concentrations. Further information on fabrication is given elsewhere [13]. Using microfabrication techniques is advantageous: it is possible to readily achieve micron-scale, or even smaller, lateral feature dimensions as well as excellent reproducibility of the structures and their behavior, since they are batch fabricated and not individually hand-made. Batch fabrication also makes the devices potentially low cost. The techniques described above are compatible with those used to make other MEMS devices, so conjugated polymers can be combined with them to form complex systems. As mentioned above, because of their low operating voltages, conjugated polymers can also be interfaced with standard integrated circuitry [7].

12.4 Single Hinge Bilayer Devices: Flaps and Lids The most commonly employed type of actuator on the microscale is the bilayer, typically attached to a single rigid plate. The bilayer comprises the polymer and a metal film. Since the device is actuated electrochemically, a nonreactive noble metal is usually used, typically gold or platinum. The electrode serves both mechanical and electrical functions. It is the non-volume-changing layer that converts inplane strain in the conjugated polymer

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to a bending motion. At the same time, it provides the electrical connection to the conjugated polymer so that it can be switched. The rigid components can be either surface-micromachined polymeric plates, as shown in Figure 12.1 and 12.2, or they can be bulk micromachined out of the substrate itself. 12.4.1

Bilayer Actuators

Because of the uniform film thicknesses, microfabricated bilayers bend into arcs of circles. The longer the bilayer, the greater the bending angle; if they are long enough, they will bend into spirals [20]. Even though the in-plane strain for PPy(DBS) is only 3%, a microfabricated PPy(DBS)/Au bilayer 30 mm long can have a radius of curvature of only 6 mm. The radius of curvature (R) is determined by the Young’s moduli of the two materials and the thickness ratio of the layers. It can be approximated using the Timoshenko equation [21]:   1  6mnð1 þ mÞ  ¼ ¼ R tp 1 þ 4mn þ 6m2 n þ 4m3 n þ m4 n2  6ð1 þ m Þ 2   ¼ 1 ðtp þ tm Þ 2 2 3ð1 þ m Þ þ ð1 þ mnÞ m þ mn

ð12:1Þ

where m is the ratio of the polymer and metal film thicknesses (m ¼ tp/tm), n is the ratio of their Young’s moduli (n ¼ Ep/Em) and a is the free actuation strain of the polymer in the plane of the film due to electrochemical switching, that is the relative change in length (D‘/‘) of an unconstrained polymer film (not attached to the metal) under no load. The curvature peaks at a certain thickness ratio. For EP ¼ 0.2 GPa and Em ¼ 83 GPa, the maximum curvature is obtained when the polymer is five times thicker than the metal [22]. To obtain more accurate predictions of the bending, it is necessary to take into account the fact that the polymer film usually does not have a uniform actuation strain in the thickness direction, but that it is higher at the free surface [19]: polymer films without a metal backing will bend by themselves. Such modeling is described elsewhere [19]. It is important to note that the free actuation strain (a) is not equal to the bending strain found from the curvature. The extent to which the bilayer is actually bent depends, as shown in Equation (12.1), on the relative thickness and stiffness of the materials. In the limit of an infinitely stiff substrate, the bilayer does not bend at all, despite the fact that a free film of the conjugated polymer would expand by several %. The strain at the outer edge of the polymer in a bilayer is not of interest in trying to determine a. The force exerted by the bilayer is proportional to the polymer film thickness and the width of the bilayer, and is inversely proportional to its length. However, increasing the thickness reduces the speed and the bending angle. Because of the various trade-offs, design of the actuator requires consideration of all of the performance requirements. 12.4.2

Drug Delivery

A device using a bilayer as a valve to open an aperture has been demonstrated by Madou et al. [23–26]. The aim is to use the device for controlled drug release, and the concept is illustrated in Figure 12.3. When the lid is in the flat, closed position, the drug is confined

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255

Figure 12.3 Illustration of a bilayer microvalve array for drug delivery. Under each flap is a vial containing the substance to be released. Actuating the bilayer causes it to lift up, opening the container (Reproduced with permission from [26]. Copyright (2006) Elsevier).

within a vial in the substrate. Upon actuation, the flap is raised, releasing the drug into the surrounding fluid. Prototypes that are able to release a fluorescent dye have been shown [26]. The release of multiple dyes in pulses was achieved by controlling the activation of an array of valves. The pulsing, or turning the supply of dye on and off, is possible because the actuation is reversible: the flaps go back to the closed position when the voltage is switched. 12.4.3

Cell Manipulation

When placed inside a microchannel, a bilayer can be used as a raisable barrier to block the passage of objects. If a good seal can be achieved between the bilayer and the channel walls, then it can also be used to control the flow of fluid [28]. Park et al. [29] used PPy flaps to control the movement of mouse embryo cells inside a microfluidic cell processor system, as illustrated in Figure 12.4; automated manipulation of

microchannel walls suction hole rotation valve (closed)

target cell

processing chamber

Figure 12.4 Overhead schematic of a cell processor for automated single cell manipulation. A PPy/Au bilayer valve was fabricated within a microchannel and used to control the passage of cells into a processing chamber (Reproduced with permission from [29]. Copyright (2005) Royal Society of Chemistry).

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individual cells would have applications in individual-cell-based diagnosis, gene injection, and in vitro fertilization. The cells were moved from a reservoir to an injection port by pressure driven flow. After a target cell had passed over the flap in the down position and moved into the processing chamber, the flap was raised to prevent other cells from entering the chamber. Inside the chamber, the cell was moved to the center by suction from a microscale hole in the substrate, and electro-rotated to control the cell orientation. After manipulation, the cell was moved to a collection well, and then the remaining cells in the reservoir were treated one by one. 12.4.4

Cell-Based Sensors

Our research group is employing bilayer actuators with rigid plates as lids to cover microvials in cell-based sensors (Figure 12.5) [30–34]. These microsystems use integrated circuits to monitor the living cells, using them as transduction elements. For example, we are developing an odorant sensor (bioelectronic nose) that uses olfactory sensory neurons to directly transduce odorant binding to receptors on the cell membrane into electrical signals, which can be detected by the IC chip. Electronic noses are of interest in a wide

Figure 12.5 Microstructures for cell based sensing. Upper left: schematic illustration of lidded vials fabricated on the surface of an integrated circuit chip with sensors on the bottom of each vial and cells cultured within them. Upper right: close-up of the bilayer actuator and the rigid plate that rotate to close the vial. Lower left: Photograph of such a system, before the final etch release. The lines visible under the microstructures are components of the sensors and circuitry. Lower right: a lidded vial actuated in cell medium in the open position with a single cell inside the vial, and in the closed position. (Fabrication and testing performed by Christophersen, M.)

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range of applications, from monitoring food freshness to detecting explosives to search and rescue. Biological odorant sensors have unparalleled sensitivity and selectivity, which we hope to be able to exploit with this approach. Beyond this application, one can also envision using such systems for rapid screening, monitoring cells for their responses to drugs and pathogens, as well as for basic biology studies. Such sensors would collect data continuously for weeks. This cell-based sensing technology is based on the combination of conjugated polymers with complementary metal oxide semiconductor (CMOS) circuitry and MEMS. It takes advantage of the ability to drive the electrochemical reactions at low voltage [7] and the ability to microfabricate the devices using CMOS-compatible processing. In the bioelectronic nose, the vials will isolate populations of neurons that display different odorant receptors on their membrane surfaces, so that the electrical signals coming from the various vials can be associated with particular molecular features. The lids will be used during the cell-loading process to ensure that only the desired vial receives a particular cell type. If these cells are very motile when cultured, as neurons tend to be, the lids will also be needed to prevent them from crawling out of the vials. Confinement by the lids is also anticipated to be necessary in other applications, such as those using nonadherent cells from the immune system. In order for the bilayers to function in cell medium at 37 °C, and in order for them to occupy minimum ‘real estate’ on the surface of the chip so that the maximum number of devices can be realized per unit area, it was necessary to carry out considerable basic characterization. For example, at body temperature the bending angle is different from that at room temperature [35], so the length of the bilayer hinges had to be adjusted. The design also had to consider the mixture of ions in cell medium: the polymer has a different affinity for different ions, and they travel within the polymer at different speeds [36], affecting the strain.

12.5 Multi-Bilayer Devices: Positioning Tools More complex devices can be achieved using multiple bilayers. As with the combination of muscle and bone, adding rigid elements to flexible bilayer hinges expands the manipulation possibilities even further, as first illustrated with a self-folding box [20] and later with additional devices. 12.5.1

Microtools

A three-finger grasping device, developed by Micromuscle AB [37], is illustrated in Figure 12.6. The opposing bilayers were surface micromachined on a substrate, and when actuated were able to grasp a small rod. As with the other devices presented here, this one operated immersed within a liquid electrolyte. 12.5.2

Microrobot

The first, and thus far still the only, microscale robot arm (Figure 12.7) was realized using a series of PPy/Au hinges and rigid plates [38–41]. It was able to pick up and move a 100 mm

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Figure 12.6 Overhead view of a grasping tool comprising three opposing bilayers. The fingers close to grasp a rod (Reprinted with permission from [37]. Copyright (2007) Elsevier).

finger actuators PP y

wrist actuator BCB elbow actuator

Au

Figure 12.7 Overhead illustration of a microrobot with multiple PPy/Au bilayer actuators and rigid plates. The sequence on the right shows actuation of the elbow and wrist to raise the arm and position it over a bead, then actuation of the fingers to grasp the bead, and actuation of the wrist and elbow to raise the bead [41].

glass bead. This application illustrates the advantages of robust, high-strain, electricallycontrolled, microfabricatable polymer actuators. No other microtechnology has been able to achieve this level of sophistication in manipulation.

12.6 Swelling Film Devices: Valves As mentioned previously, devices can also be based on out-of-plane volume expansion. This strain can be 10 times larger than the inplane strain, depending on the polymer. This is

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the case in PPy(DBS), for example, which has a lamellar structure [42] with the lamellae parallel to the electrode and ions inserted preferentially between lamellae. 12.6.1

Out-of-Plane Actuation Strain

If the out-of-plane strain is large, it can be used directly, without the need for further amplification. One advantage of this approach is that the simpler mechanical configuration results in greater actuator robustness. Another is that linear actuators produce greater forces than bending actuators. If the polymer film is kept relatively thin, then speed is not sacrificed. 12.6.2

Microvalve

A clever microvalve that opens and closes a microchannel based on swelling was demonstrated by Berdichevsky et al. [43] (Figure 12.8). The configuration consisted of a PPy(DBS) film deposited over a gold post within an electrolyte-filled chamber, which also contained a ring-shaped counter electrode and a gold pseudo-reference electrode. A 7 mm deep microchannel, containing a biological or chemical analyte, crossed over the top of the PPy-covered post. The chamber and the microchannel were formed from an elastomer (polydimethylsiloxane, PDMS), with the ceiling membrane of the actuation chamber serving as the floor of the microchannel. Upon actuation, the PPy swelled, pushing against the elastomeric membrane and causing it to move up into the microchannel, pinching it off. Because the PPy was in an isolated compartment, there were no constraints on the fluid in the microchannel based on the requirements of the PPy, and the electrolyte in the pumping chamber could be separately optimized for actuation. This allows a great deal of flexibility. (Designs with the polymer directly in the fluid path have been less successful [18].)

microchannel

Au

electrolyte

PPy

Open

PDMS

Closed

Figure 12.8 Schematic of a microvalve based on out-of-plane volume expansion. Left: the valve is open, allowing flow in the microchannel above the PPy. Right: the PPy expands, pinching off the microchannel (Reprinted with permission from [43]. Copyright (2003) Materials Research Society).

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12.7 Lifetime A question that must be addressed with any actuator is that of device lifetime: how many cycles before failure, and what are the main causes of failure? The primary cause of failure in conjugated polymer microdevices (and also macroscale devices) is delamination of the polymer from the electrode caused by extreme stresses at the interface upon cycling [22]. Delamination can be eliminated or mitigated by roughening the metal surface [44–46] and by good design (for example using a post instead of a flat surface [43] or introducing a buffer layer [22]). (Various chemical surface treatments have also been investigated, but these have either not been robust to electrochemical cycling or insufficiently reproducible.) The other cause of failure is nucleophilic attack of the polymer under electrochemical cycling in certain electrolytes. For example, in aqueous solutions the electroactivity of PPy(DBS) steadily declines, until at 35 000 cycles it is reduced to half its initial value [44]. On the other hand, in ionic liquids PPy can be cycled indefinitely [47]. Plating of the electrode onto the polymer can also be an issue in some electrolytes [48]. Another issue upon extended cycling under load is creep [12], although a significant fraction of this is recoverable. Of course, if the ultimate tensile strength is exceeded, then the actuator will rupture. The ultimate tensile strength and the Young’s modulus are functions of the polymer and dopant, the solvent and the salt in which the actuator is cycled, and the oxidation level of the polymer.

12.8 Integrated Systems As mentioned previously, one of the most important and exciting advantages of conjugated polymer microactuators is the possibility of integrating them into more complex microsystems, such as the cell-based sensors described above. The actuators add the powerful capability of mechanical manipulation of biology on the microscale. This is enabled by their low voltage operation and room temperature microfabrication, and the wide variety of actuator configurations that are possible. System integration is challenging, however. Each component must function not only individually but also all together, and the fabrication of each component must be compatible with that of all the others. The actuators must be specially designed to work under the conditions that the device will encounter, without either being damaged by or damaging their surroundings (for example through rapid swings in ion concentration or pH). To achieve a truly miniature system, integrated circuitry for controlling the actuators must also be included [7]. The process of realizing such systems consists of development cycles that start with sensor and circuit design and the fabrication of the chips in a CMOS foundry; that goes on to gold plating of the exposed aluminum (used in CMOS, but which is neither electrochemically stable nor biocompatible), fabrication of the microstructures and microactuators on the chip surface, wire bonding and packaging; and that ends with testing in the biological environment [30]. This requires a team of investigators working in close collaboration and combining different areas of expertise.

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12.9 Conclusions A variety of conjugated polymer microactuators for biomedical applications have been demonstrated in the laboratory. They have proven themselves to be practical for creating mechanical motion at the microscale and in biomedical environments, and for integration into larger systems. Recall that the only microrobot arm that has so far ever been made used conjugated polymer actuators. This puts them at the forefront of increasingly small, increasingly powerful microdevices that have the potential to begin fighting difficult medical conditions. It is now time to start considering the move from the creation of prototypes to the testing of these devices in clinical trials and consideration of initial commercialization. One company, Micromuscle AB in Sweden, has already been formed to commercialize PPy microactuators. Enough work has been done in this field to show proof-of-concept, to suggest ball-park performance and to indicate the types of designs and applications that are appropriate. The technology is mature enough to be applied in research labs, as demonstrated by the cell-manipulation system. Discussions with biologists and medical researchers could be one avenue leading to relatively rapid practical employment of the actuators. Such collaborations would offer the added benefit of experience with real-life device development and testing, without requiring substantial funding. One of the hurdles facing those who seek to commercialize the technology is the perception that ‘polymer actuators just don’t work’, based on wildly unrealistic promises that have been made by some in the EAP field in the past. Another is the erroneous but widespread perception that everything is already known about these actuators and that no further research is required. It must be appreciated that with electroactive polymers, as with any other technology, a thorough and honest understanding of the actuation behavior and performance metrics is required to achieve a well-working, consistent system, and that this takes time and investment. These myths, and the desire for amazing new technologies on the cheap, impede funding for taking these devices to the next level. Another significant challenge is the length of time it takes to gain regulatory approval for a biomedical device. The survival of the company during this extended process must be considered: what other streams of income will keep it going? The next step is to identify viable commercial applications for which these actuators are uniquely suited; like any other new technology, they should not attempt to compete with established technologies that are already in place, unless the new technology offers substantially superior performance. Only if the potential pay-off is large enough will it be possible to obtain the required investment. To identify these opportunities, those in the EAP field need to begin discussions with medical doctors, device manufacturers, and others outside the field.

References 1. Kladitis, P. E. and Bright, V. M. (2000) Prototype microrobots for micro-positioning and microunmanned vehicles, Sens. Act. A, 80 (2), 132–7. 2. Kim, C.-J., Pisano, A. P. and Muller, R. S. (1992) Silicon-processed overhanging microgripper, J. Microelectromech. Syst, 1 (1), 31–6.

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3. Smela, E., Kallenbach, M. and Holdenried, J. (1999) Electrochemically driven polypyrrole bilayers for moving and positioning bulk micromachined silicon plates, J. Microelectromech. Syst., 8 (4), 373–83. 4. Smela, E. (2003) Conjugated polymer actuators for biomedical applications, Adv. Mat., 15 (6), 481–94. 5. Hunter, I. W. and Lafontaine, S. (1992) A comparison of muscle with artificial actuators, IEEE Solid-State Sensor and Actuator Workshop, 22–25 June 1992, Hilton Head Island, SC, 178. 6. Otero, T. F. and Rodrı´guez, J. (1993) Electrochemomechanical and electrochemopositioning devices: Artificial muscles, in Intrinsically Conducting Polymers: An Emerging Technology (NATO ASI Series) (ed. Aldissi, M.), Kluwer Academic Publishers, Dordrecht, 179–90. 7. Prakash, S. B., Urdaneta, M., Christophersen, M., et al. (2008) In situ electrochemical control of electroactive polymer films on a CMOS chip, Sens. Act. B, 129 (2), 699–704. 8. Ateh, D. D., Navsaria, H. A. and Vadgama, P. (2006) Polypyrrole-based conducting polymers and interactions with biological tissues, J. Royal Soc. Interface, 3 (11), 741–752. 9. Smela, E. and Gadegaard, N. (1999) Surprising volume change in PPy(DBS): an atomic force microscopy study, Adv. Mat., 11 (11), 953–7. 10. Smela, E. and Gadegaard, N. (2001) Volume change in polypyrrole studied by atomic force microscopy, J. Phys. Chem. B, 105 (39), 9395–405. 11. Wallace, G. G., Spinks, G. M. and Teasdale, P. R. (1997) Conductive Electroactive Polymers: Intelligent Materials Systems, Technomic Publishing Co., Lancaster. 12. Smela, E., Lu, W. and Mattes, B. R. (2005) Polyaniline Actuators, Part 1: PANI(AMPS) in HCl, Synth. Met., 151 (1), 25–42. 13. Smela, E. (1999) Microfabrication of PPy microactuators and other conjugated polymer devices, J. Micromech. Microeng., 9, 1–18. 14. Jager, E. W. H., Smela, E. and Ingana¨s, O. (2000) Microfabricating conjugated polymer actuators, Science, 290, 1540–5. 15. Smela, E., Ingana¨s, O. and Lundstro¨m, I. (1995) Differential adhesion method for microstructure release; an alternative to the sacrificial layer, Transducers ‘95/Eurosensors IX, Stockholm, Sweden, 1, 218–9. ¨ hman, O. (2000) Method for the manufacturing of 16. Smela, E., Ingana¨s, O., Lundstro¨m, I. and O micromachined structures and a micromachined structure manufactured using such method, US Patent 6,103,399. 17. Zhou, J. W. L., Chan, H. Y., To, T. K. H. et al. (2004) Polymer MEMS actuators for underwater micromanipulation, IEEE-ASME Trans. Mechatronics, 9 (2), 334–42. 18. Low, L.-M., Seetharaman, S., He, K.-Q. and Madou, M. J. (2000) Microactuators toward microvalves for responsive controlled drug delivery, Sens. Act. B, 67, 149–60. 19. Christophersen, M., Shapiro, B. and Smela, E. (2006) Characterization and modeling of PPy bilayer microactuators. Part 1: Curvature, Sens. Act. B, 115, 596–609. 20. Smela, E., Ingana¨s, O. and Lundstro¨m, I. (1995) Controlled folding of micrometer-size structures, Science, 268 (23 June), 1735–8. 21. Timoshenko, S. (1925) Analysis of bi-metal thermostats, J. Opt. Soc. Am., 11, 233–56. 22. Shapiro, B. and Smela, E. (2006) Bending actuators with maximum curvature and force and zero interfacial stress, J. Intell. Mat. Syst. Struct., 18 (2), 181–6. 23. Tsai, H.-K. A., Xu, H., Zoval, J. and Madou, M. (2005) Bi-layer polypyrrole artificial muscle valves for drug delivery systems, SPIE Smart Structures and Materials, Electroactive Polymers and Devices (EAPAD), 6–10 March 2005, San Diego, CA. 24. Tsai, H. A., Zoval, J. and Madou, M. (2005) Bi-layer artificial muscle valves for drug delivery devices, Device Research Conference Digest, 2005, Santa Barbara, CA, (DRC), 1, 129–30.

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25. Tsai, H.-K. A., Ma, K.-S., Zoval, J., et al. (2006) Packaged Au-PPy valves for drug delivery systems, Smart Structures and Materials 2006: Electroactive Polymer Actuators and Devices, San Diego, CA, (SPIE), 6168, (ed. Bar-Cohen, Y.). 26. Xu, H., Wang, C., Wang, C. L., et al. (2006) Polymer actuator valves toward controlled drug delivery application, Biosens. Bioelectron., 21 (11), 2094–9. 27. Tsai, H.-K. A., Ma, K. S., Wang, C., et al. (2007) Development of integrated protection for a miniaturized drug delivery system, Smart Mat. Struct., 16 (2), S295–9. 28. Pettersson, P. F., Jager, E. W. H. and Ingana¨s, O. (2000) Surface micromachined polymer actuators as valves in PDMS microfluidic system, IEEE-EMBS Special Topic Conference on Microtechnologies in Medicine & Biology, 12–14 October 2000, Lyon, France, 334–5. 29. Park, J., Jung, S.-H., Kim, Y.-H. (2005) 12–14 Oct. Design and fabrication of an integrated cell processor for single embryo cell manipulation, Lab Chip, 5, 91–6. 30. Smela, E., Christophersen, M., Prakash, S. B. (2007) 12–14 Oct. Integrated cell-based sensors and cell clinics utilizing conjugated polymer actuators, SPIE 14th Annual International Symposium on Smart Structures and Materials, Electroactive Polymer Actuators and Devices (EAPAD), San Diego, CA, 6524 (ed. Bar-Cohen, Y.), 0G 1–10. 31. Reeves, N., Liu Y., Nelson, N. M., et al. (2004) Integrated MEMS structures and CMOS circuitry for bioelectronic interface with single cells, IEEE International Symposium Circuits and Systems (ISCAS04), 23–26 May 2004, Vancouver, BC, Canada, 3, 673–6. 32. Ji, H., Abshire, P. A., Urdaneta, M. and Smela, E. (2005) A CMOS contact imager for monitoring cultured cells, IEEE International Symposium Circuits and Systems (ISCAS05), 23–26 May 2005, Kobe, Japan, 3491–4. 33. Prakash, S. B., Abshire, P., Urdaneta, M. and Smela, E. (2005) A CMOS capacitance sensor for cell adhesion characterization, IEEE International Symposium Circuits and Systems (ISCAS05), 23–26 May 2005, Kobe, Japan, 4, 3495–8. 34. Ji, H., Dandin, M., Smela, E. and Abshire, P. (2006) Integrated fluorescence sensing for lab-on-achip devices, IEEE/NLM Life Science Systems & Applications Workshop, 13–14 July 2006, Bethesda, MD. 35. Christophersen, M. and Smela, E. (2006) Polypyrrole/gold bilayer microactuators: response time and temperature effects, SPIE 13th Annual International Symposium on Smart Structures and Materials (EAPAD), 27 February–2 March 2006, San Diego, CA, 6168 (ed. Bar-Cohen, Y.), 61680V–61681. 36. Wang, X. and Smela, E. (2006) Cycling conjugated polymers with different cations, SPIE 13th Annual International Symposium on Smart Structures and Materials (EAPAD), 27 February–2 March 2006, San Diego, CA, 6168 (ed. Y. Bar-Cohen), 61680T–61681. 37. Wilson, S. A., Jourdain, R. P. J., Zhang, Q., et al. (2007) New materials for micro-scale sensors and actuators: An engineering review, Mat. Sci. Eng. R, 56 (1–6), 1–129. 38. Smela, E., Ingana¨s O. and Lundstro¨m, I. (1995) Micro-robot controlled by conducting polymer muscles: multi-hinged, folding actuators, International Symposium on Microsystems, Intelligent Materials, and Robots, 27–29 September 1996, Sendai, Japan (eds Tani, J. and Esashi, M.), 79–82. 39. Jager, E. W. H., Smela, E., Ingana¨s, O. and Lundstro¨m, I. (1999) Polypyrrole microactuators, Synth. Met., 102 (1–3), 1309–10. 40. Jager, E. W. H., Smela, E., Ingana¨s O. and Lundstro¨m, I. (1999) Applications of polypyrrole microactuators, Proceedings of the SPIE 6th International Symposium on Smart Structure and Materials, Electroactive Polymer Actuators and Devices (EAPAD), 1–2 March 1999, Newport Beach, CA, 3669 (ed. Bar-Cohen, Y.), 377–84. 41. Jager, E. W. H., Ingana¨s, O. and Lundstro¨m, I. (2000) Microrobots for micrometer-size objects in aqueous media: potential tools for single-cell manipulation, Science, 288 (5475), 2335–8.

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42. Wernet, W., Monkenbusch, M. and Wegner, G. (1984) A new series of conducting polymers with layered structure: Polypyrrole n-alkylsulfates and n-alkylsulfonates, Makromol. Chem., Rapid Commun., 5, 157–65. 43. Berdichevsky, Y. and Lo, Y.-H. (2003) Polymer microvalve based on anisotropic expansion of polypyrrole, Mat. Res. Soc. Symp. Fall Meeting, 1–5 December 2003, Boston, MA, 782 (eds LaVan, D. A., Ayon, A. A., Madou, M. J., et al.), A4.4.1. 44. Liu, Y., Gan, Q., Baig, S. and Smela, E. (2007) Improving PPy adhesion by surface roughening, J. Phys. Chem. C, 111 (30), 11329–38. 45. Pyo, M., Bohn, C. C., Smela, E., et al. (2003) Direct strain measurement of polypyrrole actuators controlled by the polymer/gold interface, Chem. Mat., 15 (4), 916–22. 46. Cui, X. and Martin, D. C. (2003) Fuzzy gold electrodes for lowering impedance and improving adhesion with electrodeposited conducting polymer films, Sens. Act. A, 103 (3), 384–94. 47. Lu, W., Fadeev, A. G., Qi, B. H., et al. (2002) Use of ionic liquids for pi-conjugated polymer electrochemical devices, Science, 297 (5583), 983–7. 48. Smela, E. and Mattes, B. R. (2005) Polyaniline Actuators, Part 2: PANI(AMPS) in MSA, Synth. Met., 151, 43–8.

13 Actuated Pins for Braille Displays Geoffrey M. Spinks and Gordon G. Wallace ARC Centre of Excellence for Electromaterials Science and Intelligent Polymer Research Institute, University of Wollongong, Australia

13.1 Introduction The display of information on a screen is the single biggest barrier to visually impaired people accessing information. The prevalence of computer based displays and the popularity of the internet highlight the magnitude of the problem. It is estimated that the visionimpaired population of the USA is 2.5 million – a huge number of people who are largely left out of the ‘internet revolution’ and the associated employment, educational and recreational opportunities. There is tremendous need to develop a computer screen for visually impaired people that provide the same information that visual displays give to the sighted. A refreshable, two-dimensional Braille screen offers a type of ‘graphical user interface’ for visually impaired people to use with computers. Braille is used extensively in printed material to convey textual (and graphical) information. Braille books and so on are commonly available. Braille consists of patterns of raised dots (six or eight dots make up a single letter in the Braille alphabet in North America and Europe, respectively) that are ‘read’ by touch. An electronic Braille screen would use a system of raising and lowering pins to make the different patterns needed to make words, sentences and other characters. A 2-D screen would enable more complex graphics to be displayed, compared with simple text. The complexity of visual information available in a typical computer program available to sighted persons is shown in Figure 13.1. In contrast, existing electronic Braille devices show only a single line of text and cannot show sophisticated graphics (maps, music, mathematical formulae) or include functions such as drop-down menus, navigation tools and so on.

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Biomedical Applications of Electroactive Polymer Actuators Drop-down Menus

Formatting Tools

Scroll Bars

Navigation Tools

Figure 13.1 Typical word processing screen enjoyed by sighted persons.

Existing commercial electronic Braille displays only show one line of text, so there has been much effort recently devoted to the development of electronic, refreshable 2-D Braille screens. Multiple dot Braille prototypes have been built using micromachined valves to allow pneumatic actuation of rubber membrane dots [1, 2]. Like all pneumatic systems, an external pump is required and the gas venting can be noisy. Lee and Lucyszyn [3] developed micromachined heating elements that were used to melt wax resulting in a volumetric expansion to form Braille dots. Whilst the required dot height was achieved and multiple dots could be assembled and operated to the required dimensions, the cycle time (~50 s) was far too slow for a refreshable display. Other systems have been proposed based on electromagnetic solenoids [4], shape memory alloy wires [5, 6] and electrorheological fluids [7]. While all these systems have advantages and disadvantages in terms of performance and cost, this chapter explores the technical requirements for actuators to operate Braille dots and considers the performance of polymer actuators for this application.

13.2 Requirements for the Electronic Braille Screen The close spacing required for actuated pins in an electronic Braille screen significantly limit the type of actuation mechanism that can be employed. Each of the six or eight pins making up a Braille character is 1.5 mm in diameter and is spaced 2 mm from its neighbours [3] (Figure 13.2). The screen itself is expected to consist of 1000 characters or 6000–8000 individual pins in an array of 40 characters wide and 25 characters deep (Figure 13.3). This arrangement precludes the current design of electronic Braille characters that use piezoelectric ceramic bending beams (Figure 13.4), where two parallel stacks of beams gives individual control of eight pins per cell, 5–40 characters wide and a

Actuated Pins for Braille Displays 2 mm

1.5 mm

267

Pin Spring Spring support Counter electrode CP actuator Printed circuit board and connection layer

Figure 13.2 Concept design of an electronic Braille cell using a CP electrochemical actuator.

130 mm

20cells 20cells 20cells 20cells 20cells

20 cells 20 cells 20 cells

60 mm 20 cells 20 cells

Figure 13.3 Concept design of a Braille screen that is approximately A4 in size and consists of 40 cells wide and 25 cells deep (8000 individual pins).

Figure 13.4 Existing electronic Braille screen unit. Piezoelectric beam actuators (four shown) operate each pin.

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maximum of two rows. However, the length of the beams of ~40 mm means that extra rows of characters are separated by a considerable distance. The full screen display suitable for illustrating equations, music, maps and other diagrams requires close spacing of all 1000 characters on the page. The small actuation strains produced by piezoelectric ceramics (~0.1 %) mean that they are only suitable in a bending beam arrangement where the beam tip displacement produces the vertical pin movement. Alternative actuation systems are needed to produce vertical pin movement in a closely spaced arrangement. Table 13.1

Essential and desirable characteristics of electronic Braille cells

Characteristic

Essential

Desirable

Pin displacement (vertical), mm Refresh speed, s Tactile resistance

0.5

1

0.5 10 g (0.1 N) depresses pin <0.1 mm (stiffness >1000 N/m) 106 cycles

0.2 50 g (0.5 N) depresses pin <0.1 mm (stiffness >5000 N/m) 107 cycles <20 5–40 2  2  30 Mechanically robust

Lifetime Pin operating voltage, V Operating temperature, °C Physical dimensions, mm 2  2  50 Other

Several key operating requirements for pin movement in the Braille screen can be defined. These requirements are summarized in Table 13.1 and relate mainly to pin displacement, tactile resistance and refresh rate. Other requirements such as voltage, power, operating lifetime and operating environment are also important, but are more flexible and may vary between different systems. The essential characteristics are based on standard formats for Braille and typical reading speeds of vision-impaired persons. The requirements can be translated into the more standard units used to describe actuator performances. In particular, the pin displacement and allowable length of the actuator defines a strain of 1 % (essential) and 2 % (desirable). These requirements are determined from the required pin displacement of 0.5 mm and overall height of the Braille screen (50 mm maximum). The pin must sit on a spring system that resists finger pressure and, therefore, provides tactile resistance. Existing systems are designed so that 10 g finger force (0.1 N) produces less than 0.1 mm depression in the pin height. The spring, therefore, needs a stiffness of 0.1 N/0.1 mm or 1 N/mm (1000 N/m). To produce the full 0.5 mm pin displacement, the actuator must, therefore, generate 0.5 mm  1 N/mm ¼ 0.5 N of force. The maximum cross-sectional area of the actuator is 2 mm2 (allowing space for ancillary components), so that the stress generated by the actuator is 0.25 MPa. A more desirable design would produce a force of 2.5 N or stress of 1.25 MPa.

13.3 Mechanical Analysis of Actuators Operating Against Springs To further evaluate the strain/stress requirements of a potential actuation system, it is necessary to analyse the effects of operating the actuator against a restoring spring [8].

Actuated Pins for Braille Displays

269

Pin

∆ LAB

Spring

∆ LBC

Spring support CP actuator

V = 0; E = Y

V ≠ 0; E = Y

V ≠ 0; E = Y’

(a)

(b)

(c)

Figure 13.5 Actuator arrangement suitable for a Braille pin where the tactile resistance, return mechanism and pre-tensioning of the actuator are provided by a spring; (a) rest position: pin up; (b) contraction of actuator due to application of a stimulating voltage (V); (c) additional deformation of the actuator due to the voltage-induced change in Young’s modulus (E, in this case a reduction in modulus causing an increase in actuator length: Y’
A typical Braille cell arrangement involves the actuator connected to the Braille pin, but operating against a spring (Figure 13.5). The spring provides the tactile resistance and also the restoring force needed to push the pin to the ‘up’ position (slender actuators will buckle easily and may not push the pin up). The spring can also be set so that it applies a small tensioning force to the actuator to maintain the actuator in a straight position. Initially, the actuator is assumed at rest so that the pin is in the ‘up’ position. The pre-tension caused by an external spring produces an initial tension force of fA operating on the actuator. Application of a voltage causes contraction of the actuator and deformation of the spring by an amount denoted DLAB in Figure 13.5b, so that the restoring force increases with increasing actuator contraction (force increases from fA to fB). The increasing force will cause the actuator to extend, somewhat canceling its ‘natural’ contraction. It is also known that the voltage causes a change in elastic modulus (stiffness) of the actuator so that its length under load changes by an additional amount DLBC (Figure 13.5c) and the final spring force becomes fC. Since the application of the external force and the actuator stroke occur simultaneously, the effect of the external load (from the spring) on the actuator strain cannot be neglected. The actuator strain can be estimated assuming linear elastic (Hookean) behaviour of the actuator material. The strain is first estimated for the case where the elastic modulus of the actuator material does not change when stimulated. The change in length after stimulation is: DLAB ¼ DL0 þ

ðf B

f AÞ ki

ð13:1Þ

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where DL0 is the actuation length change at zero external force and ki is the actuator material’s stiffness (YA/L0, where Y is the Young’s modulus; A is the initial cross-sectional area and L0 is the initial length). The change in the external force results from deformation of the spring, so that: fB

f A ¼ ke DLAB

ð13:2Þ

and ke is the spring stiffness. The useful actuation is then: DLAB ¼

DL0 1þr

where r ¼

ke ki

ð13:3Þ

The same analysis can be extended to include the changing internal stiffness (modulus) of the actuator material during the actuator stroke. Now the actuator strain is given by the free stroke (DL0), less the displacement caused by the increasing spring force (DLAB), and less the displacement caused by a change in modulus (DLBC): DLAC ¼ DL0

fB

fA ki



fC ki 0

fB ki



ð13:4Þ

0

where ki is the internal stiffness of the actuator after the electrical stimulus has been applied (ki0 ¼ Y 0 A/L0, where Y’ is the Young’s modulus in the final state). Expanding this expression and assuming the pretension force to be small gives: DLAC ¼

DL0 1 þ r0

where r 0 ¼

ke k0i

ð13:5Þ

The results of Equations (13.3) and (13.5) are illustrated in Figure 13.6. The calculations are based on an external spring having a stiffness of 1000 N/m and a hypothetical actuator 0.80 a) b)

Actuation (mm)

0.60

c) 0.40

0.20

0.00 1

10 100 1000 10000 External Spring Stiffness (N/m)

100000

Figure 13.6 Calculated actuation length changes for hypothetical actuator operating against restoring springs of different stiffnesses. Further details given in the text. Final actuator moduli are: a) 150 MPa; b) 50 MPa; and c) 1400 MPa.

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material that produces a free strain of 3 % (in one second) and has a Young’s modulus of 150 MPa initially, and 50 MPa after stimulation. The starting length and area are taken to be 50 mm and 1 mm2, respectively. The calculations show that the actuation decreases as the external spring stiffness increases and a smaller actuation is produced when the modulus decreases during actuation (Figure 13.6b). In this example, a suitable actuation displacement (0.5 mm) is only achieved against an external spring of 1000 N/m stiffness when the Young’s modulus does not change during actuation (Figure 13.6a). Curve (c) in Figure 13.6, shows the calculated results for a much stiffer conjugated polymer (CP) actuator (final modulus of 1400 MPa) that also gives acceptable displacement against a 1000 N/m spring. The free stroke of this stiffer actuator is much smaller (1 %) compared with the first actuator example. However, the higher stiffness negates the effect of the spring force so that higher displacements are produced in this second actuator when operated against stiffer springs. These examples show that the effect of modulus and modulus change on actuation performance cannot be neglected. The analysis given here can be used to assess the viability of various actuators for operating in an electronic Braille screen given that their free stroke and final modulus values are known.

13.4 Polypyrrole Actuators for Electronic Braille Pins Polypyrrole (PPy) actuators are a type of low voltage conjugated polymer electromechanical actuator that has been evaluated for possible use in an electronic Braille screen [9]. PPy actuation occurs through an electrochemical process whereby small voltages (<5 V) applied between two electrodes in an electrolyte cause the movement of ions and solvent between the polymer and the electrolyte [10]. This mass transfer is mainly responsible for the volume changes that occur and that produce useful deformations/forces. The actuation performance of PPy has been shown to vary depending on the type of dopant ion and electrolyte used (Table 13.2). Strains and strain rates are measured under different Table 13.2

Summary of reported performances of PPy actuators

PPy formulation

Electrolyte

Maximum Actuation Strain (%)

Peak Actuation Strain Rate (% /s)

Reference

PPy/PF6 film PPy/PF6 tube with Pt helix PPy/DBS film PPy/DBS film with compliant gold PPy/TFSI Braille cell target

TBAþPF6 in PC TBAþPF6 in PC

5 5

0.1 16

[20] [9]

NaCl in H2O NaCl in H2O

6 12

0.06 0.4

[21] [21]

LiþTFSI in PC/H2O

29 2 (for 30 mm height)

5.5 8 (for 0.2 s refresh rate)

[22]

Notes: TBAþ: tetrabutylammonium; PF6–: hexafluorophosphate; PC: propylene carbonate; DBS: dodecyl benzene sulfonate; TFSI: bis(trifluoromethylsulfonyl)imide.

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conditions, but typically the applied external force is low. Very often peak values are reported, which are difficult to interpret in relation to the requirements for a Braille cell. Early work on PPy actuators showed that they had adequate strains, but response times were too slow. Improvements in PPy actuation rates have been achieved by applying techniques to increase the rate of charge injection. Madden and co-workers [11] showed that by applying resistance compensation techniques, the voltage input to the actuator could be controlled so as to shorten the response time. Their work produced an increase in actuation rate to ~3 %/s. Further increases in strain rate were achieved in PPy actuators by incorporating a flexible metal electrode into a tubular geometry (Figure 13.7). The flexible electrode was constructed from a helix of thin (50 mm) platinum wire, such that its spring stiffness was small. This design allowed for more efficient electron transfer along quite long lengths of PPy (up to several centimetres), which otherwise has quite a high resistance. Using this design, strain rates of up to 16 %/s were recorded and during continuous operation at 1 Hz, strains of 1 % were obtained [9]. The PPy helix tube actuators were fabricated into prototype electronic Braille cells, as shown in Figure 13.7b. Since the initial Braille cell prototypes using PPy actuators were reported, there have been considerable advances made in the performance of conjugated polymer actuators. In particular, Kaneto and co-workers have reported PPy formulations that can produce strains of up to 40 % [12, 13]. Such large strains typically occur at very low voltage scan rates over tens of minutes. However, these same workers have shown that strains in excess of 10 % can be obtained within two seconds [13]. In other work, polyaniline fibres reinforced with carbon nanotubes have been shown to actuate at stresses in excess of 100 MPa [14]. These fibres are easily mass-produced using a wet-spinning process, unlike the batch synthesis of PPy films and helix tubes. These new designs have not yet been tested in Braille cells, although published data suggests that they should perform adequately. The polyaniline– carbon nanotube fibers produce strains of only ~1 %, but their high modulus means that they should perform as shown by curve (c) in Figure 13.6. Another key aspect of Braille cell performance is the operating lifetime of the actuators. It is expected that individual actuators will produce 106 cycles (and preferably 107 cycles) during their operating life. This performance would provide a reasonable lifetime of the Braille screen of several years of normal use. Early conjugated polymer actuators showed poor operating lifetimes, especially when cycled over wide potential ranges to produce large movements. The discovery of room temperature ionic liquids as electrolytes for conjugated polymer systems produced very significant increases in the operating lifetime of CP devices, including actuators [15]. In conventional electrolytes, the strain produced from a PPy actuator was found to decline continuously over the first few thousand cycles. However, the same actuator operated in an ionic liquid was found to give a much more stable response to 7000 cycles [15], and further work in our own laboratories has successfully operated PPy actuators to 500 000 cycles with little degradation in performance. While the PPy helix tube actuators provide adequate performance in electronic Braille screens, further improvements in performance are required before these systems will be accepted for full-scale production. In particular, the actuator length should be reduced so that the screen has the thickness typical of a laptop computer. To achieve this dimension, the actuator should be <20 mm in length, but still produce 0.5 mm in displacement. Figure 13.8 shows how the length of the actuator affects the performance requirements needed to achieve the required displacement. The calculations shown in Figure 13.8 were based on an

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(a)

Epoxy Connector length

60 mm

Connector length A

B

C

(b)

D

E

F

Moving Pin

Spring (5 g)

Connect to Auxiliary Electrode

~200 µm Electrolyte 20~60 mm

70 mm

125 µm The PPy Hollow Fibre Connect to Working Electrode Pt Wire 9~10 mm

Figure 13.7 (a) Assembly method for preparing PPy helix tube actuators – A: a 50 m Pt wire is wound around a 250 m Pt wire; B: the wound structure is immersed in monomer solution and connected to a power supply; C: the coated electrode is removed from the electropolymerisation cell and washed; D: the inner core Pt wire is removed; E: shorter lengths of Pt wire are re-inserted into the hollow tube to form connectors; F: the connectors are fixed to the hollow helix tube using adhesive. (b) Fully assembled prototype Braille cell. (Reprinted with permission from Synthetic Metals, High performance conducting polymer actuators utilising geometry and helical wire interconnects by Ding, J., Liu, L., Spinks, G.M., et al., 138, 3, 391–8. Copyright (2003) Elsevier).

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Free Stroke (%)

6 5 4 a)

3

b)

2

c) 1 0 0

100

200 300 Final Modulus (MPa)

400

500

Figure 13.8 Minimum free stroke required (in <0.5 sec) to meet Braille cell requirements for hypothetical actuators having different moduli (after voltage stimulation) operated against a 1000 N/m spring and for different maximum actuator lengths: a) 20 mm; b) 30 mm and c) 40 mm.

actuator of cross-sectional area of 1 mm2 and operating against an external spring of 1000 N/m stiffness. The results show that quite high free strokes are needed when the modulus is <100 MPa. (Note for Braille pins this free stroke needs to be achieved in <0.5 s). The required free stroke decreases as the modulus increases but there is little advantage in achieving moduli above 200 MPa in this example. If longer actuator lengths are allowed, the required free stroke is smaller. Unfortunately, full data are not available for recently reported high-stroke CP actuators that would enable an assessment of their suitability for the Braille screen.

13.5 Other Polymer Actuation Systems for Electronic Braille Pins Several other polymer actuators have been developed for electronic Braille displays and related tactile systems [16]. Some of these developments are briefly reviewed in this section. Ionic polymer–metal composite (IPMC) bending type actuators have been built into a multi-cell Braille display prototype [17]. The prototype consisted of 24 Braille letters, each consisting of six pins (conforming to the American Braille system). A threelayer structure was used where the actuators were sandwiched between an electronics layer (below) and a sealing thin rubber layer (above). Each Braille pin consisted of a small plastic semi-sphere attached to the end of a 4  1 mm IPMC bending actuator. A clever arrangement of IPMC actuators and manufacturing process enabled the multi-pin prototype to be produced. The entire prototype structure was only 1 mm thick and was robust and flexible. The pins were driven by voltages applied to the actuators individually and the

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display was successfully tested on visually impaired persons. Although the human subjects were able to ‘read’ the displays, the performances of the actuators were below the specifications given above. In particular, the IPMC actuators could generate a maximum force of only 0.015 N, well below the suggested 0.5 N for adequate tactile resistance. The maximum displacement (no load) of the IPMC actuators was 0.4 mm, close to the target (0.5 mm), although the displacement under an external load was not provided. Response times were several seconds and the actuators relaxed slowly when the voltage was removed. Questions related to long-term performance of the IPMC actuators were also not addressed. A novel design using an electrostrictive polymer film wrapped around a pre-stressed spring has recently been described for operating an electronic Braille pin [18]. The prestressed spring core provided the tactile resistance and ‘return mechanism’. The wrapped film of terpolymer based on poly(vinylidene fluoride) expanded upon electrical stimulation to push the Braille pin vertically. A multi-layer film structure was found necessary to generate sufficient force to deform the spring. The film thickness of each layer was 6 mm, which required 600 V for operation. The authors suggest that reducing the film thickness to 2 mm would reduce the drive voltage to 200 V, but manual handling of such thin films was not possible. The wrapped springs were 34 mm initial length and actuation produced a vertical displacement of 1 mm (3 % strain) while deforming the inner spring. The spring stiffness was 1.5 N/mm, which is suitable for tactile resistance (Table 13.1). No mention was made of the speed of operation but electrostrictive polymer actuators are typically fast. Apart from the high operating voltage (and possibly lifetime) the actuator design meets all the requirements listed in Table 13.1 for an actuator suitable for electronic Braille displays. Using a quite different design, electrostrictive elastomers have also been used to build prototype Braille cells [19]. In a bubble-type arrangement, the actuation of the elastomer produced an out-of-plane deformation that pushed a pin in the vertical direction. In this design the pre-stretched elastomer itself provided the ‘spring’ for tactile resistance. No specifications were given relating to actuation displacement or force generated. The completed cell was only 9 mm in height, which falls well within the specifications for electronic Braille displays. A prototype two-cell display of six pins each was constructed and tested on visually impaired subjects. A reasonable recognition rate was achieved from the prototype.

13.6 Summary There is a great need to develop electronic Braille screens to allow easier access for visually impaired people to computer based information that now dominates modern lifestyles. These screens require multiple actuators that can be individually addressed and positioned beneath each Braille pin. Due to the close spacing and small size of the pins, conventional motor systems are unsuitable. Actuator materials are ideal for this application and the development of the electronic Braille screen represents a tremendous opportunity to develop a worthwhile, commercial product using polymer actuators. The specifications required of an actuator material to operate Braille pins are well defined, with a displacement of >0.5 mm in <0.5 s the key requirement. Since there must be ‘tactile resistance’ to enable the pin to be ‘read’, the actuators will almost inevitably operate

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against a restoring spring. The external load applied by the spring to the actuator will change during the actuation cycle (as the spring deforms), so that the final actuator length will depend on its own stiffness. The effect of external spring force can be successfully modelled using a simple linear elastic approach and the model can be used to assess different actuator materials and different Braille cell designs. Various prototype electronic Braille cells have been constructed using polypyrrole, IPMC and electrostrictive elastomer actuator materials. In each case, some further improvements are needed to meet all the essential and desirable requirements for the electronic Braille screen.

Acknowledgements The authors thank Tim Connell and John Gillespie from Quantum Technology Pty. Ltd (Sydney, Australia) for providing information relating to Braille screen specifications and providing concept designs. John Gillespie provided Figures 13.2–4.

References 1. Yobas, L., Lisy, F.J. and Durand, D.M. (2001) A Novel Bulk-Micromachined Electrostatic Microvalve with a Curved-Compliant Structure Applicable for a Pneumatic Tactile Display, J. Microelectromech. Syst., 10, 187–96. 2. Yobas, L., Durand, D.M., Skebe, G.G., et al. (2003) A Novel Integrable Microvalve for Refreshable Braille Display System, J. Microelectromech. Syst., 12, 252–63. 3. Lee, J.S. and Lucyszyn, S. (2005) A Micromachined Refreshable Braille Cell, J. Microelectromech. Syst., 14, 673–82. 4. Frisken-Gibson, S.F., Bach-y-Rita, P., Tompkins, W.J. and Webster, J.G. (1987) A 64-Solenoid, four level fingertip search display for the blind, IEEE Trans. Man-Machine Syst., BME-34, 963–5. 5. McCarty, L.H. (1990) Special alloy is the key to Braille computer display, Design News, 2 December 1990, 158–9. 6. Taylor, P.M., Moser, A. and Creed, A. (1998) A sixty-four element tactile display using shape memory alloy wires, Displays, 18, 163–8. 7. Taylor, P.M., Pollet, D.M., Hosseini-Sianaki, A. and Varley, C.J. (1998) Advances in electrorheological fluid based tactile array, Displays, 18, 135–41. 8. Spinks, G.M. and Truong, V.-T. (2005) Work-per-cycle analysis of electromechanical actuators, Sensors and Actuators A: Physical, 119, 455–61. 9. Ding, J., Liu, L., Spinks, G.M., et al. (2003) High Performance Conducting Polymer Actuators Utilising a Tubular Geometry and Helical Wire Interconnects, Synth. Met., 138, 391–8. 10. Refer to the chapter of this book on the fundamentals of conducting polymer actuators for more details on the mechanisms of actuation. 11. Madden, J.D., Cush, R.A., Kanigan, T.S. and Hunter, I.W. (2000) Fast contracting polypyrrole actuators, Synth. Met., 113, 85–92. 12. Zama, T., Tanaka, N., Takashima, W. and Kaneto, K. (2006) Fast and large stretching bis(trifluoromethylsulfonyl)imide (TFSI)-doped polypyrrole actuators and their applications to small devices, Polym. J., 38, 669–77. 13. Hara, S., Zama, T., Takashima, W. and Kaneto, K. (2006) Tris(trifluoromethylsulfonyl)methidedoped polypyrrole as a conducting polymer actuator with large electrochemical strain, Synth. Met., 156, 351–5.

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14. Spinks, G.M., Mottaghitalab, V., Bahrami-Samani, M., et al. (2006) Carbon Nanotube Reinforced Polyaniline Fibres for High Strength Artificial Muscles, Adv. Mat., 18, 637–40. 15. Lu, W., Fadeev, A.G., Qi, B.H., et al. (2002) Use of ionic liquids for pi-conjugated polymer electrochemical devices, Science, 297, 983–7. 16. Refer to the related chapters of this book. 17. Kato, Y., Sekitani, T., Takamiya, M., et al. (2007) Sheet-Type Braille Displays by Integrating Organic Field-Effect Transistors and Polymeric Actuators, IEEE Trans.Electron Devices, 54, 202–209. 18. Ren, K., Liu, S., Lin, M., Wang, Y. and Zhang, Q.M., A compact electroactive polymer actuator suitable for refreshable Braille display, (2007) Proceedings of SPIE Smart Structures and Materials 2007: Electroactive Polymer Actuators and Devices (EAPAD) (ed. Bar-Cohen, Y.), 6524, G1–9. 19. Lee, S., Jung, K., Koo, J., et al. Braille Display Device Using Soft Actuator, (2004) Proceedings of SPIE Smart Structures and Materials 2004: Electroactive Polymer Actuators and Devices (EAPAD) (ed. Bar-Cohen, Y.) 5385, 368–79. 20. Spinks, G.M., Liu, L., Zhou, D. and Wallace, G.G. (2002) Strain response from polypyrrole actuators under load, Adv. Funct. Mat., 12, 437–40. 21. Bay, L., West, K., Sommer-Larsen, P., et al. (2003) A conducting polymer artificial muscle with 12% linear strain, Adv. Mat., 15, 310–3. 22. Hara, S., Zama, T., Takashima, W. and Kaneto, K. (2005) Free-standing gel-like polypyrrole actuators doped with bis(perfluoroalkylsulfonyl)imide exhibiting extremely large strain, Smart Mat. Struct., 14, 1501–10.

14 Nanostructured Conducting Polymer Biomaterials and Their Applications in Controlled Drug Delivery Mohammad Reza Abidian1 and David C. Martin1, 2, 3 1 Biomedical Engineering, The University of Michigan, Ann Arbor, MI, USA Materials Science and Engineering, The University of Michigan, Ann Arbor, MI, USA 3 Macromolecular Science and Engineering, The University of Michigan, Ann Arbor, MI, USA 2

14.1 Introduction The design, synthesis and development of conjugated electrically and ionically conducting polymers has opened many new opportunities to address important problems in science, technology and medicine. The p–p orbital overlap that comes from the alternating single–double bonds along the chain backbone means that these materials have extended electronic states, in some sense similar to conventional metals, but now confined to an organic molecular framework. This leads to a variety of interesting and important properties including electrical and ionic conductivity, light absorption and emission, and the ability to efficiently exchange both anions and cations with a surrounding electrolyte [1, 2]. These materials, including the relatively chemically stable functionalized thiophenes like poly(3,4-ethylenedioxythiophene) (PEDOT), are of considerable current interest for use at the abiotic–biotic interfaces in biomedical devices. They can readily accommodate the dramatic differences in properties between electrically conducting, stiff, flat, inorganic metals and semiconductors, and the ionically conducting, soft, wet, organic, living tissue. Conducting polymer coatings are currently being considered for a variety of biomedical devices, including microfabricated cortical electrodes, cochlear implants, artificial retinas,

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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pacemakers, deep brain stimulators and glucose sensors [3–6]. Here, the processing of these materials into various nanostructures and their utility as biomedical device coatings with drug delivery capabilities are described [7]. The ability to process conducting polymers into intricate, nanostructured shapes makes it possible to optimize their performance. Nanostructured materials exhibit unique and novel properties, largely as a consequence of their small size [8, 9]. They have a wide range of applications in chemistry [10, 11], physics [12], biology [13], electronics [14, 15], mechanics [16, 17] and materials science [18, 19]. They are especially important for applications in biomedical sciences such as drug delivery [7, 20], neural engineering [21, 22], biosensors [23] and biotechnology and drug discovery [24]. Nanomaterials can exist in a wide variety of geometries including nanoparticles, coreshell nanostructures, hollow nanospheres, nanofibers, nanotubes and nanocomposites. They can be created from metals, semiconductors, synthetic organic and biological materials. The widespread interest in nanostructured materials mainly originates from the fact that their properties (optical, electrical, mechanical and chemical performance) can be substantially different from those of the corresponding bulk materials. The preparation of these objects using self-assembly and templating techniques has been described in detail in several excellent reviews [25–30]. For example, porous membranes can be used to synthesize conducting polymer nanotubes [31].

14.2 Nanostructured Conducting Polymers Conducting polymers (CPs) have attracted considerable attention since the initial discovery of polyacetylene in the late 1970s [1, 2, 32, 33]. Conducting polymers have a conjugated double-bonded backbone that provides electronic conductivity after doping. Their response to electrochemical oxidation or reduction can produce a simultaneous change in conductivity [1, 34], color [35–37] and volume [38, 39]. Conducting polymers are of current interest for a wide variety of devices including sensors, actuators and displays. The reversible doping of conducting polymers, with associated control of the electrical conductivity over the full range from insulator to metal, can be accomplished either by chemical doping or by electrochemical doping. In electrochemical doping, the electrode supplies the redox charge to the conducting polymer, while ions diffuse into (or out of) the polymer structure from the nearby electrolyte to compensate the electronic charge [40]. 14.2.1

Fabrication

A variety of fabrication methods have been developed for nanostructured conducting polymers. Among the various synthetic strategies, templating is a promising and powerful tool to fabricate conducting polymer nanomaterials. Templating involves growing conducting polymers inside the void spaces of a host material. The void spaces act as the template, determining the shape, size and orientation of the compound produced [18, 41]. They can be designed in forms of nanoparticles, hollow nanoshperes, nanofibers and nanotubes. Poly(3,4-ethylenedioxythiophene) (PEDOT) has been fabricated in the shape of rods, tubes, thimbles and belts through chemical polymerization in the pores of aluminum oxide membrane [42].

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Recently, we have developed a novel method for fabrication of polypyrrole (PPy) and PEDOT nanotubes by electrochemical polymerization around electrospun polylactide (PLA) nanofiber templates [3, 7]. After PPy and PEDOT deposition, the PLA fibers were removed by soaking in dichloromethane. The wall thickness of the PEDOT nanotubes varied from 50–100 nm and the nanotube diameter ranged from 100–600 nm. By controlling the polymerization time, we could reproducibly prepare tubular structures with thin walls or thick walls. Scanning electron micrographs of PPy and PEDOT nanotubes on neural probe electrode sites are shown in Figure 14.1. The total applied charge density during electrochemical deposition was 1.44 C/cm2 (applied current 20 nA for 900 s) for all samples. The overall thickness of PPy and PEDOT nanotube coatings depends upon the layer thickness of the nanofibers and CP deposition charge density. Figure 14.2 shows surface morphologies of nanotubular PPy doped with LiClO4 electrochemically polymerized on the electrode sites as a function of applied charge density from 0.48 to 2.88 C/cm2. The current density was controlled during the deposition at 0.16 mA/cm2.

Figure 14.1 Scanning electron micrographs of electropolymerized PPy and PEDOT nanotubes on neural microelectrode sites. (A) Top view of PPy nanotubes; (B) three-dimensional view of PPy nanotubes; (C) top view PEDOT nanotubes; and (D) three-dimensional view of PEDOT nanotubes. PPy nanotubes and PEDOT nanotubes with deposition charge density 1.44 C/cm2 [3] (Reprinted with permission from Biomaterials, Experimental and theoretical characterization of implantable neural microelectrodes modified with conducting polymer nanotubes by Abidian, M. R. and Martin, D. C., 29, 9. Copyright (2008) Elsevier).

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Figure 14.2 Scanning electron micrographs of electropolymerized PPy nanotubes on neural microelectrode sites as a function of deposition charge. (A) 6 mC, (B) 9 mC, (C) 12 mC, (D) 18 mC, (E) 24 mC and (F) 36 mC. [3] (Reprinted with permission from Biomaterials, Experimental and theoretical characterization of implantable neural microelectrodes modified with conducting polymer nanotubes by Abidian, M. R. and Martin, D. C., 29, 9. Copyright (2008) Elsevier).

Several synthetic methods for preparing PEDOT nanoparticles have been reported including seed polymerization, emulsion polymerization and dispersion polymerization. There have been several reports related to PEDOT-coated particles and PEDOT hollow particles [43, 44]. Dispersion polymerization has been applied for PEDOT-coated Polystyrene (PS) particle fabrication. 100 nm PS nanoparticles were used as the core material [44]. Polyaniline (PANi) nanofibers have been synthesized using interfacial polymerization without templates or functional dopants [45, 46]. Scanning electron microscopy (SEM) images of PANi nanofibers are shown in Figure 14.3. 14.2.2

Biomedical Application

Conducting polymers are being considered for a number of organic electronic device applications such as light-emitting diodes [47, 48], thin film transistors [49–51], supercapacitors [52–56], photovoltaic cells [57–59], and chemical sensors [60–66]. Some of the most exciting and important applications are expected from biomedical devices, where they have been considered as actuators [7, 67–71] and sensors [4, 7, 67, 72–78]. Conducting polymers are well suited to biomedical applications. They are generally biocompatible both in vitro and in vivo for long periods [76, 77, 79–84]. Conducting polymers are of interest for tissue engineering because they not only physically support tissue growth, but they are also electrically and ionically conductive and thus should be able to stimulate specific cell functions or trigger cell responses [76, 85–87]. They can be doped with a variety of pharmaceutical agents and released using electrical stimulation [88–92].

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Figure 14.3 SEM images showing the morphology of polyaniline synthesized from a) a rapidly mixed reaction and b) a slowly mixed reaction. High quality nanofibers are obtained in the rapidly mixed reaction, while irregular agglomerates form in the slowly mixed reactions (Reprinted with permission from Angewandte Chemie, Nanofiber formation in the chemical polymerization of aniline: A mechanistic study by Huang, J. and Kaner, R. B., 43, 5817–21. Copyright (2004) Wiley-VCH GmbH Co. KGaA).

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Recently, Isaksson et al. demonstrated that the controlled transport of calcium and potassium cations using PEDOT based devices could be used to locally manipulate the physiology of cells in vitro [93]. Conducting polymers can be modified by adding certain chemical functional groups that can change the degree of cell adhesion and cell-surface interactions [94, 95]. Bioerodable and biodegradable conducting polymers have been produced with potential as a temporary scaffold for cell attachment and as a source of electrical signals to stimulate tissue regeneration [81, 82]. Conjugated polymer actuators are already being used for blood vessel connectors, Braille displays and cochlear implants [96]. Other devices still in the research stage include microvalves, steerable catheters, microvials for single-cell studies and drug delivery [7, 69, 97]. Many of the studies to date have involved nerve, bone, muscle or cardiac cells that respond to electrical impulses [98]. Conducting polymers present a number of important advantages for biomedical applications [76], including the ability to deposit on the surface of electrodes and transfer charge from ions in living tissue to electrons in an electrode [93], and the ability to entrap and controllably release biological molecules [90]. These characteristics are useful for applications such as biosensors [73], tissue-engineering scaffolds, neural microelectrodes [99], drug delivery devices [7] and bio-actuators [67]. Among the known conducting polymers, we have been particularly interested in PPy, PEDOT and PEDOT derivatives because of their promising electrical properties and biocompatibility [76, 79]. In addition to its high conductivity (circa 300 S/cm), PEDOT has been found to be almost transparent in thin film and shows high chemical stability in the oxidized state [35, 36, 100]. We have found that soft, low impedance and biologically active coatings can be prepared by the electrochemical deposition of these conducting polymers on neural microelectrode arrays [3, 4, 6, 101]. Neural microelectrodes are designed to provide control signals for neural prosthetic devices via transduction of biological signals to electronic signals [102, 103]. Studies have shown only a few recording neural microelectrodes continue to function for long periods [103–105]. Cellular reactive responses that are considered to contribute to device failure include an early acute inflammatory response due to insertion trauma and a chronic foreign body reaction induced by tethering, micromotion and device biocompatibility [106–108]. These responses are characterized by activation of glial cells and eventually formation of a nonconductive sheath layer referred to as ‘glial scar’, which causes a progressive increase in electrode-tissue interface impedance [106–112]. Therefore, a low impedance electrode-tissue interface is important for maintaining signal quality for recording as well as effective charge transfer for stimulation. However, neural microelectrode sites typically exhibit high impedances because of their small geometric surface area. Several studies have been done to explore strategies to reduce the initial impedance electrode sites or to limit early and late reactive responses or in order to minimize electrode-tissue interface impedance, including electrochemical deposition of conducting polymers on the surface of recording sites [3, 7, 101, 113] and systematic release of antiinflammatory drugs to the site of implantation [109, 114]. Electrochemically deposited films of PEDOT and PPy make it possible to provide better contact and a larger surface between the electrode and neuronal tissue [4, 6, 7, 99, 113, 115–119]. Scanning electron micrographs of the electrode sites reveal polymer growing on the gold electrodes, creating a fuzzy surface with a large surface

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area to maximize neuron–electrode interactions. It was found that the impedance of coated sites significantly decreased and the maximum reduction of impedance was observed for PEDOT nanotubes [7]. More recently we found that surfactanttemplated ordered PEDOT coated electrodes had lower impedance and a higher quality of neural recording (with signal-to-noise ratios increased by about 17 %) in vivo out to six weeks [84].

14.3 Conducting Polymer Nanotubes for Controlled Drug Delivery During the past decade, polymeric microsphere and hydrogel materials have been shown to be effective in enhancing drug targeting, lowering systematic drug toxicity and improving treatment absorption rates. Many controlled release systems have been designed to improve the effectiveness of drugs using hydrogels [120, 121] and biodegradable polymers [20, 122]. The normal difficulties of delivering drugs to the central nervous system are compounded with the difficulties in delivering bioactive molecules over long implantation periods. Current delivery methods include PPy/polypeptide blends grown on electrode pads [4, 115], electrical stimulation of conducting polymers containing bioactive molecules [7, 88, 90] and sustained release of dexamethasone from polymeric matrices [110, 123]. Shain et al. found that peripheral injections of dexamethasone at the time of electrode insertion greatly attenuated glial scar formation at one and six weeks as shown by GFAP staining. Some attenuation was seen with local release of dexamethasone from implanted poly(ethylvinyl) acetate strips, but at six weeks post-implantation the effect was minor [110]. Wells etched into polyimide electrode developed by Rousche et al. (2001) were filled with dextran as a proof of concept, but could potentially hold other diffusible compounds or hydrogels. Chen et al. (1997) attempted to address this drug delivery problem with the development of ‘puffer’ probes that incorporate microfluidic channels inside the electrode shank. These bulk machined silicon probes have multiple 10 mm channels for chemical and drug delivery from orifices typically situated 2.5 mm from recording sites [124]. Conducting polymers have the ability to actuate using external electrical stimulation [68, 71], which makes it possible to design controlled drug delivery systems [7, 93]. Their response to electrochemical oxidation or reduction can produce a change in volume. When ions and solvent enter the polymer it expands and when they exit the polymer it contracts [70, 90, 125]. This makes it possible for the conducting polymer to locally deliver a drug by a change in oxidation state, as directed by an externally applied potential (Figure 14.4). Such a device is able to administer therapeutic agents such as neurotrophic proteins [88, 90, 126] and anti-inflammatory drugs [7, 127] at controllable dosages that provide optimum therapeutic effects and prevent undesirable side effects. Conducting polymers can be used directly as carriers for drug molecules that are ionically coupled to the bulk of conducting polymer either as dopants during electropolymerization or as charged ions during reduction and oxidation. We recently reported on a method to prepare conducting-polymer nanotubes that can be used for precisely-controlled drug release [7]. As shown in Figure 14.5, the fabrication process involves the electrospinning of a biodegradable polymer (here poly(lactide-co-glycolide)

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Reduction + e–

Negatively charged counterion or drug

Figure 14.4 Schematic illustration showing controlled drug release from a drug-incorporated conducting polymers using a reduction process.

(PLGA)), into which a drug (here dexamethasone, an anti-inflammatory drug) has been incorporated, followed by electrochemical deposition of a conducting polymer (here poly(3,4-ethylenedioxythiophene) PEDOT) around the drug-loaded, electrospun biodegradable polymers. Dexamethasone can be released from the PEDOT nanotubes in a desired fashion by subsequent electrical stimulation of the PEDOT nanotubes; this process presumably proceeds by a local dilation or contraction of the tube that promotes mass transport. 14.3.1

Electrospinning

Electrospinning, a drawing process based on electrostatic interactions, provides a simple approach to creating polymer nanofibers with both solid and hollow interiors that are exceptionally long in length, uniform in diameter and diverse in composition [128–131]. Unlike other methods for generating one-dimensional nanostructures, the formation of a thin fiber via electrospinning is based on the uniaxial stretching (or elongation) of a viscoelastic jet derived from a polymer solution. Electrostatic spinning or electrospinning forms fibers two to three orders of magnitude smaller in diameter than conventionally processed fibers [130, 132]. Electrospinning occurs when a charged solution or melt of a high molar mass polymer is subjected to an electric field. Because of the presence of chain entanglements in the charged fluid, the fluid does not break up into discrete droplets but forms a stable jet when the electrostatic repulsive forces on the fluid surface overcome the surface tension. The range of fiber diameters is roughly between 100 nm and 10 mm [133] and is dependent on: (i) processing variables, such as the electrical field strength, fluid flow rate and working distance between the electrodes [134]; (ii) solution variables, such as the

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Figure 14.5 Schematic diagrams illustrating the surface modification of neural microelectrodes to create nanotubular PEDOT: (A) electrospinning of biodegradable polymer (PLGA) fibers with well-defined surface texture (1) on the probe tip; (B) electrochemical polymerization of conducting polymers (PEDOT) (2) around the electrospun fibers; and (C) dissolving the electrospun core fibers to create nanotubular conducting polymers (3) [7] (Reprinted wih permission from Advanced Materials, Conducting polymer nanotubes for controlled drug release by Abidian, M. R., et al., 18, 4, 405–9. Copyright (2006) Wiley-VCH Verlag GmbH Co. KGaA). (See Color Plate 5).

viscosity, electrical conductivity, surface tension and solvent volatility [135]; and (iii) environmental variables, including temperature, pressure and humidity [136, 137]. Figure 14.6 shows a schematic illustration of the basic setup for electrospinning. It consists of three major components: a high-voltage power supply, a spinnerette (a metallic syringe needle) and a collector (a grounded conductor). 14.3.2

Electrospinning of Dexamethasone-Loaded Template PLGA Nanofibers

Poly(lactide-co-glycolide) (PLGA) is suitable for the template since it can be readily processed into nanofibers, is stable during electrochemical deposition of the conducting polymer coating and can be easily removed under conditions that leave the wall material intact. A blend of PLGA and dexamethasone was dissolved in chloroform to create a solution of PLGA and dexamethasone. Then, dexamethasone-loaded PLGA fibers were

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Figure 14.6 Typical electrospinning process setup [140] (Reprinted with permission from Journal of Applied Physics, Bending instability of electrically charged liquid jets of polymer solutions in electrospinning by Reneker, D. H., Yarin, A. L., Fong, H. and Koombhongse, S., 87, 9, 4531–47. Copyright (2000) American Institute of Physics).

collected on the conductive substrate though electrospinning process. The diameter of drug-loaded electrospun fibers typically ranged from 40 to 500 nm with the majority being in the range 100–200 nm. Since the chloroform was a volatile solvent, a well defined surface texture was observed on the nanofibers [138]. 14.3.3

Electrochemical Deposition of PEDOT Nanotubes

EDOT monomer was electrochemically polymerized onto the conductive electrode substrates, growing up and around the dexamethasone-loaded PLGA nanofibers using an Autolab PGSTAT 12 (EcoChemie, Utrecht, The Netherlands) in galvanostatic mode with a conventional four-electrode configuration at room temperature (working and sensing electrodes were connected together and reference and counter electrodes were connected together). An aqueous solution of EDOT monomer was prepared in phosphate buffered saline (PBS). Optical micrographs that show the various steps of the fabrication

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Figure 14.7 Optical micrographs of: (E) the gold electrode site; (F) the electrode site after electrospinning showing the coverage of the PLGA electrospun nanoscale fibers; (G) the electrode after electrochemical deposition of PEDOT on the gold site and around the electrospun fibers; and (H) the electrode after removal of the core nanoscale fiber templates (Reprinted with permission from Advanced Materials, Conducting polymer nanotubes for controlled drug release by Abidian, M. R., et al., 18, 4, 405–9. Copyright (2006) Wiley-VCH Verlag GmbH Co. KGaA). (See Color Plate 6).

process are presented in Figure 14.7: conductive substrate before electrospinning (E); dexamethasone-loaded PLGA nanofibers collected on substrate (F); electrochemical deposition of PEDOT around the electrospun drug-loaded nanofibers (G); PEDOT nanotubes after degradation of PLGA nanofibers (H). An FEI Focused Ion Beam (FIB) instrument was used to make cross-sections of the PEDOT nanotubes; as shown in Figures 14.8C and 14.8D, the PEDOT nanofibers are hollow. Figures 14.8E and 14.F show the hollow PEDOT nanofiber morphology with well-defined, rough internal and external texture. The internal texture is replicated from the external texture of PLGA electrospun nanofibers [139]. The wall thickness of the PEDOT nanotubes varied from 50 to 100 nm and the nanotube diameter ranged from 100 to 600 nm. The nanotubular structures could be prepared with different wall thickness by controlling the polymerization time. 14.3.4

Controlled Drug Delivery from PEDOT Nanotubes

In order to detect the release of dexamethasone from the electrospun PLGA nanofibers and the PEDOT nanotubes, UV spectrophotometry was used. The absorption was monitored at a wavelength of 237-nm, which corresponds to the absorbance peak of dexamethasone. The amount of incorporated dexamethasone within PLGA nanoscale fibers on each sample was approximately 2 mg. PLGA nanofibers had a high ratio of surface area to volume, therefore, their degradation happened quickly due to fast hydrolytic degradation of electrospun nanofibers. It was found that about 75 % of the dexamethasone loaded in PLGA nanofibers was released after seven days; that was a burst effect. Coating the PEDOT layer around the drug-loaded electrospun PLGA nanofibers dramatically slowed down the rate of drug release. Although the burst

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Figure 14.8 Scanning electron micrographs of PLGA nanoscale fibers and PEDOT nanotubes. (A) diameters of the PLGA fibers were distributed over the range 40–500 nm with the majority being between 100 and 200 nm. (B) Electropolymerized PEDOT nanotubes on the electrode site of an acute neural probe tip after removing the PLGA core fibers. (C) A section of (B) cut with a FIB showing the silicon substrate layer and PEDOT nanoscale fiber coating. (D) Higher magnification image of (C) showing the PEDOT nanotubes crossing each other. (E) A single PEDOT nanotube which was polymerized around a PLGA nanoscale fiber, followed by dissolution of the PLGA core fiber. This image shows the external texture at the surface of the nanotube. (F) Higher-magnification image of a single PEDOT nanotube demonstrating the textured morphology that has been directly replicated from the external surface of the electrospun PLGA fiber templates. The average wall thickness of PEDOT nanotubes varied from 50 to 100 nm, with the nanotube diameters ranging from 100 to 600 nm (Reprinted with permission from Advanced Materials, Conducting polymer nanotubes for controlled drug release by Abidian, M. R., et al., 18, 4, 405–9. Copyright (2006) Wiley-VCH Verlag GmbH Co. KGaA).

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effect was completely eliminated from the dexamethasone release profile, less than 25 % of the dexamethasone was released after 54 days, presumably because the diffusion of dexamethasone through the walls of the PEDOT nanotubes was difficult (dexamethasone was not water soluble), requiring the molecules to move to the tube ends before being released. On the other hand, a small portion of drug diffused into the solution in the absence of any PEDOT actuation, while the reminder was trapped inside the PEDOT nanotubes. Since it was reported that conducting polymers could be actuated using electrical stimulation [68, 70, 125], we hypothesized that the PEDOT nanotubes could deliver dexamethasone by applying a bias voltage and thus actuating the drug-loaded PEDOT nanotubes. We successfully demonstrated controlled release of the drug by electrical stimulation of the PEDOT nanotubes. To electrochemically control the nanotube actuation, an Autolab A

C

B

D

E

F CE

CE

v

v

WE

WE

Electrolyte

Dexamethasone Anion Cation

CE Counter Electrode WE Working Electrode

Figure 14.9 Schematic illustration of the controlled release of dexamethasone: (A) dexamethasone-loaded electrospun PLGA; (B) hydrolytic degradation of PLGA fibers leading to release of the drug; and (C) electrochemical deposition of PEDOT around the dexamethasone-loaded electrospun PLGA fiber slows down the release of dexamethasone (D); (E) PEDOT nanotubes in a neutral electrical condition; (F) external electrical stimulation controls the release of dexamethasone from the PEDOT nanotubes due to contraction or expansion of the PEDOT (Reprinted with permission from Advanced Materials, Conducting polymer nanotubes for controlled drug release by Abidian, M. R., et al., 18, 4, 405–9. Copyright (2006) Wiley-VCH Verlag GmbH Co. KGaA). (See Color Plate 7).

Biomedical Applications of Electroactive Polymer Actuators 2.0

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Cumulative Mass Released (mg)

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1.5 1.0 0.5 0.0

0

200 400 600 800 1000 1200 1400 Time (h) (a)

0.8 0.6 0.4 0.2 0.0

220 240 260 280 300 Wavelength (nm)

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(b)

Figure 14.10 (a) Cumulative mass release of dexamethasone from: PLGA nanoscale fibers (black squares), PEDOT-coated PLGA nanoscale fibers (red circles) without electrical stimulation, and PEDOT-coated PLGA nanoscale fibers with electrical stimulation of 1 V applied at the five specific times indicated by the circled data points (blue triangles). (b) UV absorption of dexamethasone- loaded PEDOT nanotubes after 16 h (black), 87 h (red), 160 h (blue) and 730 h (green). The UV spectra of dexamethasone have peaks at a wavelength of 237 nm. Data are shown with a – standard deviation (n ¼ 15 for each case) (Reprinted with permission from Advanced Materials, Conducting polymer nanotubes for controlled drug release by Abidian, M. R., et al., 18, 4, 405–9. Copyright (2006) Wiley-VCH Verlag GmbH Co. KGaA). (See Color Plate 8).

PGSTAT 12 galvanostat/potentiostat was used with a conventional four-electrode configuration. A platinum wire was the counter electrode and Ag/AgCl was the reference electrode. The drug-loaded PEDOT nanotubes were actuated by applying a positive voltage of 1 V with a scan rate of 0.1 V/s for 10 s at five specific times. As illustrated in Figure 14.9F, during reduction of the PEDOT nanotubes (positive voltage bias), electrons were injected into the chains and positive charges in the polymer chains were compensated. To maintain overall charge neutrality, negatively charged counterions were presumably expelled towards the solution and the nanotubes contracted. PEDOT contraction then produced a mechanical force creating pressure within the nanotubes. We hypothesized that the hydrodynamic force inside the nanotubes caused expulsion of PLGA degradation products and dexamethasone, probably either through the ends of PEDOT nanotubes or through openings or cracks on the surface of the nanotubes created by actuation. After electrical excitation a significant increase in the amount of dexamethasone released was observed (Figure 14.10). The externally applied voltage bias provided a means of controlling the release of the drug, as has been seen for other conducting polymers [90]. Here, the expansion and contraction of the nanotube provides an additional means of controlling the kinetics of drug release. It was anticipated that the amount of drug release would be directly related to the contraction force and the duration of the contraction, as controlled by the externally applied voltage. The fact that the drugs are released from within the nanotube chamber presumably means that much more of the active agent can be incorporated into these films than in previous systems where the pharmaceutical is embedded within the conducting polymer directly.

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14.4 Conclusions Conducting polymers are of interest for biomedical applications because of their interesting properties such as ionic conductivity, actuation, and incorporation of drugs. Templating is a promising method for fabrication of conducting polymer nanomaterials that may enhance the physical, chemical, and bioactivity properties relevant to living tissue compared to that in the bulk state. We have described methods to prepare conducting polymer nanotubes using templated electrochemical growth around nanoscale electrospun filaments. The ability to actuate the conducting polymers with an external field makes it possible to create films that can locally deliver drugs such as anti-inflammatory agents and neurotrophins. The external bias makes it possible to squeeze the encapsulated drugs out of the nanotubes in a controlled manner. We believe this method provides a generally useful means for creating low impedance, biologically active polymer coatings, which will facilitate integration of electronically active devices with living tissues. Other potential biomedical applications of these molecule-eluting, electrically active polymer nanotubes include highly localized stimulation of neurite outgrowth and guidance for neural tissue regeneration, and spatially and temporally controlled drug delivery for ablation of specific cell populations.

Acknowledgements The authors express their appreciation to Professor Daryl Kipke and The Center for Neural Communication Technology, which is a P41 Resource Center funded by the National Institute of Biomedical Imaging and Bioengineering (NIBIB, P41 EB002030) and supported by the National Institutes of Health (NIH). This research was supported in part by the National Institute of Health (NIH-NINDS-N01-NS-1-2338), the National Science Foundation (DMR-0084304), the Army MURI on Bio-Integrating Structural and Neural Prosthetic Materials, Grant W911NF-06-1-0218, the University of Michigan College of Engineering Translational Research (GAP) program and the National Academies Keck Futures Initiative. Mohammad Reza Abidian thanks the Rackham Graduate School for support through a Doctoral Fellowship. David Martin is a Founder and Chief Scientific Officer for Biotectix LLC, a company that is actively investigating the potential of conducting polymer coatings for use in biomedical device applications.

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79. George, P. M., Lyckman, A. W., LaVan, D. A., et al. (2005) Fabrication and biocompatibility of polypyrrole implants suitable for neural prosthetics. Biomat. 26, 3511–9. 80. Wang, X. D., Gu, X. S., Yuan, C. W., et al. Evaluation of biocompatibility of polypyrrole in vitro and in vivo. J. Biomed. Mat. Res. Part A 68A, 411–22 (2004). 81. Rivers, T. J., Hudson, T. W. and Schmidt, C. E. (2002) Synthesis of a novel, biodegradable electrically conducting polymer for biomedical applications. Adv. Funct. Mat. 12, 33–7. 82. Zelikin, A. N., Lynn, D. M., Farhadi, J., et al. (2002) Erodible conducting polymers for potential biomedical applications. Angew. Chem., Int. Ed. 41, 141–4. 83. Richardson-Burns, S. M., Hendricks, J. L. and Martin, D. C.. Polymerization of the conducting polymer poly(3,4-ethylenedioxythiophene) (pedot) around living neural cells. Biomat., 28, 8, 1539–52. 84. Ludwig, K. A., Uram, J. D., Yang, J. Y., et al. (2006) Chronic neural recordings using silicon microelectrode arrays electrochemically deposited with a poly(3,4-ethylenedioxythiophene) (pedot) film. J. Neural Eng. 3, 59–70. 85. Aoki, T., Tanino, M., Sanui, K., et al. (1996) Secretory function of adrenal chromaffin cells cultured on polypyrrole films. Biomat. 17, 1971–4. 86. Collier, J. H., Camp, J. P., Hudson, T. W. and Schmidt, C. E. (2000) Synthesis and characterization of polypyrrole-hyaluronic acid composite biomaterials for tissue engineering applications. J. Biomed. Mat. Res. 50, 574–84. 87. Ateh, D. D., Vadgama, P. and Navsaria, H. A. (2006) Culture of human keratinocytes on polypyrrole-based conducting polymers. Tissue Eng. 12, 645–55. 88. George, P. M., LaVan, D. A., Burdick, J. A., et al. (2006) Electrically controlled drug delivery from biotin-doped conductive polypyrrole. Adv. Mat. 18, 577–81. 89. Entezami, A. A. and Massoumi, B. (2006) Artificial muscles, biosensors and drug delivery systems based on conducting polymers: A review. Iranian Polym. J. 15, 13–30. 90. Pernaut, J. M. and Reynolds, J. R. (2000) Use of conducting electroactive polymers for drug delivery and sensing of bioactive molecules. A redox chemistry approach. J. Phys. Chem. B 104, 4080–90. 91. Cosnier, S., Senillou, A., Gratzel, M., et al. (1999) A glucose biosensor based on enzyme entrapment within polypyrrole films electrodeposited on mesoporous titanium dioxide. J. Electroanal. Chem. 469, 176–81. 92. John, R., Spencer, M., Wallace, G. G. and Smyth, M. R. (1991) Development of a polypyrrolebased human serum-albumin sensor. Analytica Chimica Acta 249, 381–5. 93. Isaksson, J., Kjall, P., Nilsson, D., et al. (2007) Electronic control of ca2þ signalling in neuronal cells using an organic electronic ion pump. Nature Ma. 6, 673–9. 94. Lee, J. W., Serna, F., Nickels, J. and Schmidt, C. E. (2006) Carboxylic acid-functionalized conductive polypyrrole as a bioactive platform for cell adhesion. Biomacromolecules 7, 1692–5. 95. Sanghvi, A. B., Miller, K. P. H., Belcher, A. M. and Schmidt, C. E. (2005) Biomaterials functionalization using a novel peptide that selectively binds to a conducting polymer. Nature Mat. 4, 496–502. 96. Immerstrand, C., Holmgren-Peterson, K., Magnusson, K. E., et al. (2002) Conjugated-polymer micro- and milliactuators for biological applications. Mrs Bulletin 27, 461–4. 97. Jager, E. W. H., Immerstrand, C., Peterson, K. H., et al. (2002) The cell clinic: Closable microvials for single cell studies. Biomed. Microdevices 4, 177–87. 98. Guimard, N. K., Gomez, N. and Schmidt, C. E. (2007) Conducting polymers in biomedical engineering. Prog. Polym. Sci. 32, 876–921. 99. Cui, X. Y., Hetke, J. F., Wiler, J. A., et al. (2001) Electrochemical deposition and characterization of conducting polymer polypyrrole/pss on multichannel neural probes. Sensors and Actuators A: Physical 93, 8–18.

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15 Integrated Oral Drug Delivery System with Valve Based on Polypyrrole Thorsten Go¨ttsche and Stefan Haeberle Institut fu¨r Mikro- und Informationstechnik of the Hahn-Schickard-Gesellschaft (HSG-IMIT), Germany

15.1 Introduction IntelliDrug is the name of an intelligent, controllable drug delivery system, proposed and patented by Dr. Andy Wolff and Eng Ben Z. Beiski of Saliwell Ltd. Israel, that can be inserted into the oral cavity of an adult human, for example as part of a dental prosthesis. Within the development of the device, two major innovations in the field of portable drug delivery systems have been addressed [1–6]. A novel route for drug delivery via the buccal mucosa has been evaluated on the one hand, and a controllable and miniaturized drug delivery system with the size of two teeth has been molar designed and fabricated on the other hand. The first innovation, the investigation of a novel route for the delivery of active agents into the body through the buccal mucosa, is illustrated in Figure 15.1. Established implantable drug delivery systems apply drugs subcutaneously, to the adipose tissue, the vascular system or via a catheter directly to the spinal cord [7]. These concepts imply considerable costs and surgical risks. The IntelliDrug system combines the advantages of implantable and extracorporeal drug delivery systems. Being part of a denture, the system allows a predetermined, controlled and discrete drug delivery without the need of a mechanical penetration of the body and the associated risks. Drug concentration in blood plasma, expected upon discrete delivery events (solid curve), is explained in Figure 15.2. If the concentration is not within a specific therapeutic window (TW), poor efficacy or undesired side effects of the drug will occur. The

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Figure 15.1 Illustration of a novel route for drug delivery: from a tooth implant to the buccal tissue. plasma concentration

risk of side effects

therapeutic window (TW)

poor efficacy

time

Figure 15.2 Therapeutic window in drug therapy.

dashed curve shows the drug plasma concentration following a more sophisticated drug administration, as aspired by modern drug delivery systems [8–11]. Here, the drug can achieve the highest therapeutic effect without the risk of side effects. The IntelliDrug system represents such an advanced drug delivery system, enabling the delivery of drugs to the body via new pathways with increased bioavailability and reduced patient compliance problems. The treatment of addiction with Naltrexone was selected as a pilot application for the IntelliDrug system. Therefore, the bioavailability of the suggested route for drug delivery through the buccal mucosa was investigated in a series of in vivo experiments in the first phase of the project. The levels of Naltrexone in blood plasma from 264 drawn blood samples, originating from 22 individual experiments with six female pigs, were measured (Figure 15.3). In 14 experiments, a solution of 10 mg of active agent with 80 mg/ml was delivered within 10 minutes directly to the buccal mucosa (this corresponds to 125 ml solution at 12.5 ml/min). In eight experiments, 10 mg (2 ml at a concentration of 5 mg/ml) were administered intravenously (i.v., bioavailability 100 %) for reasons of comparison. The graphs in Figures 15.2 and 15.3 show the eminent applicability of delivering drugs via the buccal mucosa to the systemic circulation, leading to elongated plasma levels within the therapeutic window compared to i.v. administration. The second innovation of the IntelliDrug system, namely the integration of a fully functioning and controllable microdosage system the size of two molar teeth, is the main focus of this chapter. Control of the system is accomplished using an osmotic pump, a flowand fill-level sensor [6] and a valve to control the flow rate of the drug out of the system. The programmability, addressed by IR communication through the cheek, allows the

concentration in blood plasma [ng/mL]

[ng/mL]

Integrated Oral Drug Delivery System with Valve Based on Polypyrrole

50

i. v. delivery

40

buccal delivery

30 20

140,0

10

120,0

0 30

100,0

303

90

150 210 270 time [minutes]

330

80,0 60,0 40,0 20,0 0,0 0

60

120 180 240 time [minutes]

300

360

Figure 15.3 Blood plasma levels of Naltrexone after buccal and intravenous (i.v.) delivery of 10 mg of the drug.

Table 15.1

Boundary conditions of the IntelliDrug system

• harsh environment (wet, corrosive) • temperature range 0 60 C • load up to 250 N when chewing • small volume (corresponding to two buccal teeth in an adult)

• drug capacity for 2 weeks • energy supply for 2 weeks • refillability • insertability and exchangeability without surgery

delivery of drugs according to the individual needs of the patient. The resulting technological challenges are the minimal space requirements and, corresponding to that, the need for high-level integration. In order to achieve reasonable times of usage before replacement or maintenance of the device is required, the energy consumption over two weeks of operation must stay below the capacity of two button cells. Application-specific boundary conditions for the system are listed in Table 15.1. Based on these demands, polypyrrole (PPy) was chosen as actuator material and has been investigated in different microvalve configurations as described in the course of this chapter.

15.2 System Concept Today, there is no drug delivery system with the same degree of miniaturization and complexity as the IntelliDrug system. Due to the limited space, the device is designed

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Figure 15.4 Schematic illustration of the IntelliDrug system.

to deliver a maximum amount of drug solution requiring a minimum amount of energy. As the IntelliDrug system will be located in a wet environment and quick regulation sequences are not required, osmosis offers a simple and efficient solution for flow and pressure generation [2]. The basic concept of the liquid actuation is depicted in Figure 15.4. The drug is stored as a solid pill in the drug reservoir, maximizing the drug capacity of the device. Water from the saliva enters the system through an osmotic membrane, dissolves the drug pill and thus generates an osmotic pressure in the drug reservoir. A polymer balloon, filled with air, allows ‘storage’ of pressurized drug solution that can be released by opening a microvalve. For medical safety reasons, this valve needs to be designed as ‘normally closed’ to prevent unwanted delivery of drug solution when the valve is not actuated. To keep the energy consumption of the device and the required actuation voltages at a minimum, polypyrrole (PPy), an ionic Electroactive Polymer (EAP), was selected as the actuator material for the valve. The use of a flow sensor in combination with a fill-level sensor [6] allows the monitoring of the flow rate to the buccal mucosa and the detection of an empty reservoir. Figure 15.5 shows an exploded view of the total device: the cover with the osmotic membrane (a) closes the drug reservoir in the main housing (b); the valve, the sensors and a polydimethylsiloxane (PDMS) flow channel connecting these two components are accessible for maintenance by the buccal cover (c). The volume of the drug pill, measuring 96 mm3, and the energy supply with two button cells of 1.5 V each allow application of the device for at least two weeks before an exchange of the batteries, the osmotic membrane and the drug pill is required. The amount of drug delivered is controlled by measuring the flow rate of the drug solution through the outlet during the times that the valve is open.

Integrated Oral Drug Delivery System with Valve Based on Polypyrrole

main housing lingual cover with osmotic membrane

305

valve

compressible fluidic capacity

buccal cover PDMS sealing with channel

solid drug pill

batteries

(a)

folded 4-layer PCB with support

(b)

(c)

Figure 15.5 Exploded illustration of the IntelliDrug system.

15.3 Osmotic Pressure Pump The drug delivery device is characterized by two modes of operation: ‘valve closed’, that is no consumption of electrical energy and an increase of the osmotic pressure; and ‘valve open’, that is consumption of energy for opening the valve and a decay of the pressure that is ‘stored’ in the drug reservoir (air-filled balloon). While an estimation of these two modes is given in the following section, more information on osmotic systems is given elsewhere [12–14]. 15.3.1

Valve Closed

In this mode, water from the saliva enters the device with flow rate Jw. It dissolves the drug, raises the relative pressure (pc) inside the drug reservoir from ambient pressure (p0) to a maximal achievable osmotic pressure (Dp) and compresses a fluidic capacity (air-filled balloon) from volume Vc0 at ambient conditions to volume Vc, while no drug solution is delivered: Jw þ

dVc ¼0 dt

ð15:1Þ

At constant temperature: pc Vc ¼ p0 Vc0 ¼ const:

ð15:2Þ

Combining and differentiating Equations (15.1) and (15.2) gives: Jw ¼

dVc 1 dpc ¼ p0 Vc0 dt pc2 dt

ð15:3Þ

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With definition of the osmotic flow according Equation (15.4) (compare [13]), the transient osmotic pressure generation inside the device is obtained according Equation (15.5): Jw ¼ KAðDp

ð15:4Þ

DpÞ

dpc KA ððDp þ p0 Þp2c ¼ p0 Vc0 dt

p3c Þ

ð15:5Þ

pressure pc

800

flow rate Jw

600

0,4 0,3

400

0,2

200

0,1

0 0

50

100 150 time [min]

200

0 250

flow rate Jw [µl/h]

pressure pc [kPa]

where K is the water permeability of the membrane, A is the area of the membrane,  is the rejection rate of the membrane (ideal semipermeability:  ¼ 1), Dp is the pressure difference across the membrane (¼ pc) and Dp is the difference of the osmotic pressures of the solutes at both sides of the membrane. With A ¼ 20 mm2,  ¼ 1 and K ¼ 4  10 7 ml/Pa h mm2 and Vc0 ¼ 10 mm3), Jw and p(t) are given in Figure 15.6.

Figure 15.6 Closed valve: water flow rate (Jw) into the IntelliDrug system results in an increase in pressure (pc).

15.3.2

Valve Open

With the fluidic resistance (R) of the drug delivery channel and the gas law according Equation (15.6), the decrease in pressure inside the drug reservoir can be calculated as given in Equation (15.7): dVc d p0 Vc0 1 dpc Þ ¼ p0 Vc0 ¼ ð dt pc pc2 dt dt dpc KA 1 Þp0 Þpc2 ððDp þ ð1 þ ¼ p0 Vc0 KAR dt

ð1 þ

ð15:6Þ 1 Þp 3 Þ KAR c

ð15:7Þ

The transient behaviour of the pressure pc inside the drug reservoir upon opening of the valve and the delivery of drug with flow rate Jd to the buccal mucosa according Equation (15.8) are given in Figure 15.7 (R ¼ 7.2  10þ15 Pas/m3). Jd ¼

pc

p0 R

ð15:8Þ

Integrated Oral Drug Delivery System with Valve Based on Polypyrrole

pressure pc

600

150

flow rate Jd

400

100

200

50

0

flow rate Jd [µl/h]

200

800 pressure pc [kPa]

307

0 0

2

4 6 time [min]

8

10

Figure 15.7 Open valve: flow rate of drug solution (Jd) and resulting decrease in pressure (pc) inside the drug reservoir.

The results show that the pressure that is build up in the chamber at closed valve within 4 hours is released in around 10 minutes upon opening of the valve.

15.4 Polypyrrole in Actuator Applications 15.4.1

Why PPy in the Intellidrug System

The technical goal of a intra-oral drug delivery system is quite challenging, for which reason established system components and functionalities should be resorted to whenever possible. Consequently, piezo actuators have been considered first for the microvalve. While their energy consumption as well as their size could be adapted to the IntelliDrug system, the limited space capacity prohibits the integration of an adequate voltage transformation for driving the piezo actuator with 3 V from two button cells. As a result, from the actuator point of IntelliDrug, new grounds had to be broken. Here, Electroactive Polymers (EAPs) offer great potential [15, 16]. The main actuator criteria of PPy (negatively doped with DBS–) are summarized in Table 15.2. Table 15.2

Actuator relevant information on PPy

electrical control with low voltages (1 3 V) high expansion (out of plane <25 %, lateral <5 %) suitable for microfabrication (photolithographic patterning) suitable for actuation in aqueous electrolytes containing small mobile ions (e.g. physiological saline solution) þ aqueous actuation voltage directly correlates deformation/deflection actuation times (<1s 100 s) necessity for surrounding appropriate electrolyte solution little information on actuation forces available limited information on actuation parameters available little information on long-term characteristics available

þ þ þ þ

Summarizing the information given in Table 15.2, PPy can be identified as an actuator material which geometric deformations are up to 200 times higher than the ones known from ceramic piezo actuators. A main benefit is the low actuation voltage, below 3 V

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Biomedical Applications of Electroactive Polymer Actuators

(available from the two button cells in the IntelliDrug system). While the requirement of a surrounding actuation electrolyte needs to be addressed by design, the relatively long actuation time plays a minor role in the intended application. Unfortunately, limited information on realistic actuation forces, detailed actuation parameters and long-term characteristics is currently available. Up-to-date, the only commercial provider of PPy for actuator applications is Micromuscle AB, Sweden [17]. As a close cooperation partner in the IntelliDrug project, Micromuscle AB provided the polypyrrole samples for the presented evaluations and kindly supported the research regarding actuation parameters and geometric design. 15.4.2 15.4.2.1

Actuation of PPy Setup of Actuation

The established standard for the actuation of ionic (conjugated) EAPs is a setup of three electrodes in combination with a potentiostat. All electrodes are immersed into a common electrolyte solution. The potential between a working electrode (WE ¼ the EAP actuator) and a counter electrode (CE) is regulated in that way, that the potential between the WE and a reference electrode (RE, usually Ag/AgCl) corresponds to a desired value. Due to the limited available space, it is not desired to incorporate potentiostat circuitry inside the IntelliDrug device. Consequently, the required potentials for actuating PPy with only two electrodes (WE-CE) had to be evaluated. Figure 15.8 shows the potentials WE-RE and WE-CE when actuating a sample of PPy (gold wire 225 mm in diameter, coated with PPy to a diameter of 350 mm in dry state, length 10 mm) in a 0.1 M NaDBS electrolyte (CE ¼ stainless steel, RE ¼ Ag/AgCl). Despite the small potential span RE-WE of only 0 to 1 V, potentials of þ0.7 to 1.5 V need to applied between the PPy sample and the used CE when omitting the RE, as aimed for the intended application. 1,2E–03

current [A]

8,0E–04 4,0E–04 0,0E+00 –4,0E–04 WE-CE

–8,0E–04

WE-RE

–1,2E–03 –2

–1,5

–1

–0,5 voltage [V]

0

0,5

1

Figure 15.8 Cyclic voltammograms, related to the potentials WE-RE and WE-CE.

15.4.2.2

Selection of the Electrolyte

In a series of experiments, a variety of electrolytes were evaluated. Suitable electrolytes for the actuation of PPy were found to be watery solutions of NaDBS, NaCl, Na2SO4, LiCl in the range of concentration from 0.01 to 0.5 M (physiological NaCl solution corresponds to a concentration of 0.0154 M) and artificial saliva. Solutions of Naltrexone are not suited for

Integrated Oral Drug Delivery System with Valve Based on Polypyrrole

309

the actuation of PPy. Chemical dissolution of the PPy was observed after overnight storage in a saturated solution of the drug. Therefore, in the current application, a separation between the electrolyte used for actuation and the liquid to be delivered itself is required. Liþ has a larger hydration shell than Naþ and thus drags more water into the polymer as Naþ. As water contributes to the expansion, this helps to explain that the total achievable deformations are slightly larger when actuating PPy in 0.1 M LiCl compared to 0.1 M NaDBS or 0.1 M NaCl. Since solutions containing chloride tend to show higher corrosion and the aim is for a small number of involved components (DBS– is already incorporated in the applied PPy), watery solutions of NaDBS were selected as electrolytes in the IntelliDrug system. Best actuator performance was obtained in the concentration range 0.1–0.2 M. 15.4.2.3

Selection of the Counter Electrode

In miniaturized technical applications, especially when the volume of the electrolyte is limited and encapsulated, the formation of gas by hydrolysis is a critical issue. Hydrolysis occurs at potentials in an electrochemical cell higher than 1.24 V. Therefore, an appropriate selection of the material for the counter electrode (CE) needs to be addressed when designing a miniaturized ionic EAP system. Voltammograms of a PPy/Au bilayer (area: 10 mm2 – 2 mm2, thickness: 30 mm) in 0.1 M NaDBS are given in Figure 15.9. Each plot refers to a different electrode material applied as CE: gold on chromium on silicon,

Figure 15.9

Influence of the CE material on the actuation voltages of PPy.

platinum, stainless steel and silver. In each case, the sample was actuated in a three-electrode setup by linearly ramping the potential WE(PPy)-RE(Ag/AgCl) between 0 and 1 V within 200 s. The potentials that are applied between PPy and CE differ significantly. In the case of a CE fabricated in silver, a full redox cycle could be obtained with WE-CE potentials between 0.5 and 1.2 V (compare the potentials for the oxidation and reduction peaks in current). In this case, the potentials can be kept below the threshold voltage for the occurrence of hydrolysis. Furthermore, the applied voltages do not need to be switched from positive to negative values, which simplifies required control electronics.

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15.5 Valve Concepts Evaluated in the Course of the IntelliDrug Project 15.5.1

Wafer-Level Fabricated Membrane Valve

Due to the incompatibility of the targeted drug solution with PPy, separation of the driving electrolyte and the flow path was required. For reasons of safety in medical drug delivery and optimization in energy consumption, the presented membrane valve was designed to be normally closed. Due to the strong demands on miniaturization of the IntelliDrug system, PPy membrane valves were developed and realized on wafer-level by means of silicon micromachining. 15.5.1.1

Technological Realization

Figure 15.10 shows the basic setup: a silicon wafer was oxidized and anisotropically etched from the back to realize pores 50 mm in diameter (distance 150 mm) that are closed by a thin membrane of oxide. A layer of chromium and gold was deposited and structured on the front of the wafer, serving as electrode for a local electrochemical deposition of PPy. A ring of chromium, circular located around the pores in the silicon, serves as adhesion layer for the gold. Applying a subsequent dry etching process from the back of the wafer allows a removal of the oxide membranes and thus an opening of the pores towards the gold layer. An electrolyte reservoir, closed by a flexible membrane, allows an actuation of the valve independent to its environment. Applying a potential to the on-chip CE results in a buckling of the PPy/Au bilayer and, consequently, an opening of the fluidic passage between the two pores. The gold membrane, facing towards the pores in the silicon, provides a separation between the PPy and the controlled liquid and, furthermore, prevents ionic exchange between the driving electrolyte and the drug solution. The photograph in Figure 15.10 shows the valve chip with the PPy/Au bilayer membrane (¼WE), measuring 200 mm in diameter and 10 mm in thickness and the on-chip realized CE and RE.

Figure 15.10 Wafer-level fabricated membrane valve. Left: photograph of the valve chip. Right: schematic cross-section and materials.

15.5.1.2

Measurement

Figure 15.11 shows a typical fluidic measurement, revealing two major drawbacks of the described configuration. Applying an exemplary pressure drop of 150 kPa across the valve, a reproducible manipulation of the flow between 0.6 and 1.3 ml/min was achieved upon

Integrated Oral Drug Delivery System with Valve Based on Polypyrrole 2,0

2,0 detachment of the membrane

1,0

1,5 1,0

0,5

0,5 actuation voltage

flow

0,0

0,0

0,5

–0,5

1,0

–1,0

–1,5

–1,5 0

1000

2000

pH2O = pNaDBS = 150kPa

3000 4000 time [s]

5000

6000

flow [µl/min]

actuation voltage [V]

1,5

–2,0

311

–2.0 7000

Figure 15.11 Fluidic characterization of the wafer-level fabricated PPy membrane valve. Pressure drop across the valve: 150 kPa.

application of a potential to the PPy of 1.6 V vs CE (Au). Despite these promising results, a complete tight closing of the valve could not be achieved. According to fluidic calculations, the flow rates in the not-actuated and actuated correspond to a buckling of the PPy/Au bilayer membrane of 0.9 mm and 1.3 mm upon actuation of the PPy. An explanation for the gap in the not-actuated state can be found in an initial swelling of the PPy upon its first contact to a wet environment (e.g. electrolyte solution). A further challenge is to achieve a more robust adhesion of the membrane to the silicon substrate. The flow rate and photograph in Figure 15.11 beyond an actuation time of 4200 s reveal a detachment of the PPy/Au bilayer from the valve substrate. 15.5.2

Micro-Assembled Membrane Valve

An alternative setup of a membrane valve was designed to overcome the difficulties with the initial swelling of the PPy upon first contact with the electrolyte. The described approach allows a pre-conditioning of the PPy/Au bilayer membranes prior to assembly. The applied membrane measures 1.5 mm in diameter and 20 mm in thickness. The silicon components of the valve (depicted in Figures 15.12 and 15.13) are fabricated by means of laser structuring in combination with a wet etching process. To allow the membrane to buckle and thus to give way to the drug solution to the outlet even at osmotic pressures of up to 600 kPa, both sides of the PPy membrane are pressure compensated, except the outlet opening which measures 100 mm in diameter. Therefore, a freestanding, flexible membrane of parylene, measuring 5 mm in thickness, separates the driving electrolyte from the osmotic drug reservoir. A silicone septum, cast with Sylgard 186 (Dow Corning), allows the filling of the electrolyte reservoir using cannulas. The high fluidic resistance of the flow sensor downstream to the valve prevents a collapse of the pressure at the outlet when the valve is open and thus maintains the pressure compensation.

312

Biomedical Applications of Electroactive Polymer Actuators from drug reservoir silicone septum

CE

inlet

Parylene membrane - electrolyte reservoir PPy membrane Si Si to flow sensor outlet

silicone channel and sealing valve seat

Figure 15.12 principle.

Schematic cross-section of the PPy membrane valve, illustrating its working

Figure 15.13 PPy valve with electrolyte reservoir for valve actuation in diverse environments, e.g. air or drug solution.

15.5.2.1

Technological Realization of the Freestanding Parylene Membrane

The free standing parylene-C membrane, measuring 5 mm in thickness, was realized by means of a ‘substractive’ process (Figure 15.14). The silver-electrolyte reservoir (a) was inserted into silicone tubing (b) and filled with hot, liquid wax (c). After hardening of the wax and removal of the silicone tubing (d), the silicone septum was cast into a recess in

Integrated Oral Drug Delivery System with Valve Based on Polypyrrole

313

Figure 15.14 Fabrication process for freestanding Parylene-C membranes for pressure compensated encapsulation of the actuation electrolyte.

the reservoir housing (e). Subsequent to a Parylene-C-coating (f), the wax was released by immersion in acetone, resulting in an electrolyte reservoir including a silicone septum, homogeneously covered and sealed by a partly free standing membrane of Parylene-C (g). 15.5.2.2

Measurement

Actuation, and thus buckling, of three actuator membranes is shown in Figure 15.15. The complete valve consists of an actuator membrane between two silicon components as depicted in Figure 15.12. To ease assembly, the round actuation membranes are processed with a rigid, surrounding frame. According to a spectrophotometric surface scan, a deflection of the PPy membranes upon actuation in electrolyte solution of 200 – 50 mm was achieved (Figure 15.16). The experimental flow characteristics of a microvalve with a counter electrode which was in this phase of evaluation fabricated in biocompatible stainless steel (Sandvik Bioline 316LVM) are shown in Figure 15.17. At an applied differential pressure across the valve of 50 kPa and an actuation voltage of 2.5 V vs CE, a delivery of 0.9 ml/min at an energy consumption of < 10 mJ per actuation was realized, which meets the specifications. The decrease in flow rate with an increasing number of actuation cycles is not clarified yet. Further evaluations are needed

Figure 15.15 Structured PPy/Au elements immersed in electrolyte solution. Left: no actuation. Right: buckling of the flexible round membrane (Ø1.5 mm) upon actuation.

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Biomedical Applications of Electroactive Polymer Actuators

1.5 mm

Figure 15.16 Optical measurement results of a PPy/Au bilayer membrane. Left: 0 V vs CE, Right: 2.5 V vs CE (stainless steel).

1,2

flow [µl/min]

1 0,8 0,6 0,4 0,2 0 1800

2300

2800 time [s]

3300

Figure 15.17 Fluidic characterization of the micro-assembled PPy membrane valve. Pressure drop across the valve: 50 kPa.

regarding long-term stability (so far: <50 cycles), repeatability and possible degradation effects of the polymer.

15.6 Total Assembly and Clinical Testing of the IntelliDrug System The strongly miniaturized integration of the control electronics and the electrical connection of the sensors and the valve is achieved by a four-layer flex-rigid printed circuit board (PCB), provided by Valtronic SA, Switzerland. A polymer battery housing, attached to the PCB, allows the integration of two button cells. Figure 15.18 shows the PCB assembly prior to folding and insertion into the IntelliDrug system. The reversible integration of the IntelliDrug system into a denture is depicted in Figure 15.19. The basic therapeutic applicability was proven in a clinical human study at the end of 2007 using a simplified system without valve. Despite the function of the PPy valve being proven in the first experiments, a clinical application will need further evaluation regarding reliability and long-term behaviour.

Integrated Oral Drug Delivery System with Valve Based on Polypyrrole

315

flow sensor

recesses for valve

recesses for batteries

Figure 15.18

Figure 15.19

Assembled PCB with battery housing prior to folding.

IntelliDrug system ready assembled and inserted in a model of a human denture.

Acknowledgement The IntelliDrug project was supported by a European Grant under the Sixth Framework – Project IntelliDrug – IST-FP6 Contract No 002243. The overall project management is done by Assuta Hospital (Israel). All partners of the IntelliDrug consortium are: Assuta Medical Centers (Israel), Institut fuer Biomedizinische Technik (IBMT) of Fraunhofer Gesellschaft zur Foerderung der angewandten Forschung e. V. (Germany), Relsoft Systems Ltd. (Israel), Institut fuer Mikro-und Informationstechnik (HSG-IMIT) of Hahn-Schickard-Gesellschaft (Germany), Valtronic S.A. (Switzerland), Warsaw University of Technology (Poland), Universita di Palermo (Italy), Biodar Ltd. (Israel), Promedt Consulting GmbH (Germany),

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Biomedical Applications of Electroactive Polymer Actuators

ASM Centrum Badan i Analiz Rynku (Poland), Hospital Clinico San Carlos de Madrid (IMSALUD, Spain), Faculty of Medicine and Surgery of University ‘Federico II’ of Naples (Italy), Universitaets Klinikum Charite´ (Germany), Anti Drug Authority (ADA, Israel), Telefo´nica (Spain), Saliwell Ltd. (Israel and Germany).

References 1. Project homepage of the IntelliDrug project: www.intellidrug.org. 2. Goettsche, T., Schumacher, A., Zengerle R., et al. (2007) Highly Integrated Oral Drug Delivery System As Functional Tooth, Proceedings of the Smart Systems Integration Conference, 27–28 March 2007, Paris, France, 83–90. 3. Goettsche, T., Schumacher, A., Zengerle, R., et al. (2007) Highly Integrated Oral Drug Delivery System With Valve Based On Electro-Active-Polymer, Conference on Micro-Electro-Mechanical Systems 2007 (MEMS07), 21–25 January 2007, Kobe, Japan, 461–4. 4. Scholz, O., Wolff A., Schumacher, A., et al. (2007) Drug delivery from the oral cavity: focus on a novel mechatronic delivery device, Drug Discovery Today, doi:10.1016/j.drudis.2007.10.018. 5. Go¨ttsche, T. and Wolff, A. (2006) IntelliDrug – An Integrated Intelligent Oral Drug Delivery System, mstnews, 1/06, 36–7. 6. Velten, T., Schuck, S., Knoll, T., et al. (2006) Intelligent intraoral drug delivery microsystem, Proceedings of the Institution of Mechanical Engineers, Part C, J. Mech. Eng. Sci., 220, 1609–17. 7. Ochs, G. (2004) Die Behandlung der schweren Spastizita¨t, 2nd edn, Georg Thieme Verlag, Stuttgart, New York. 8. Mathiowitz, E. (1999) Encyclopedia of Controlled Drug Delivery, John Wiley & Sons Ltd, Chichester. 9. D’Aquino, R., et al. (2004) Good Drug Therapy: It’s Not Just The Molecule – It’s The Delivery, CEP-magazine, 100, 2, 15S–17S. 10. Okano, T., Bae, Y.H. and Kim, S.W. (1990) Pulsed and self-regulated drug delivery, 17–45, CRC Press, Boca Raton, FL. 11. Razzacki, S., et al. (2004) Integrated microsystems for controlled drug delivery, Adv. Drug Del. Rev., 56, 185–98. 12. Baker, R.W. (2000) Membrane Technology and Applications, John Wiley & Sons Ltd, Chichester. 13. Theeuwes, F., et al. (1975) Elementary Osmotic Pump, J. Pharm. Sci., 64, 12, 1987–91. 14. Theeuwes, F., et al. (1976) Principles of the design and operation of generic osmotic pumps for the delivery of semisolid or liquid drug formulations, Ann. Biomed. Eng., 4, 4, 343–53. 15. Jager, E. W. H., Smela, E. and Ingana¨s, O. (2000) Microfabricating conjugated polymer actuators, Science, 290:5496, 1540–5. 16. Berdichevsky, Y., et al. (2004) Polymer Based on Anisotropic Expansion of Polypyrrole, Mat. Res. Soc. Symp. Proc., 782, A4.4. 1–7. 17. Company homepage Micromuscle AB, Sweden: www.micromuscle.com, accessed 2008.

Section IV Piezoelectric and Electrostrictive Polymers

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

16 Piezoelectric and Electrostrictive Polymer Actuators: Fundamentals Zhimin Li1 and Zhongyang Cheng2 1 2

Pharmaceutical Chemistry, Auburn University, Cambridge, Massachusetts, USA Materials Research and Education Center, Alkermes Inc., Auburn, Alabama, USA

16.1 Introduction The electromechanical (E-M) effect, which can convert energy between the electric form and mechanical form, is the foundation for many current technologies and plays a critical role in the development of the next generation. Although there are different mechanisms to create the E-M effect, the piezoelectric and electrostrictive effects have been extensively studied. Piezoelectric crystals and ceramics are the primary materials for current EM devices, such as sonar transducers, pressure/stress sensors, ultrasonic transducers, high frequency filters, acoustic resonators, ultrasonic motors, actuators and microelectromechanical systems (MEMS). From an E-M device point of view, polymeric materials provide many unique advantages, including high flexibility, light weight, high stress impact resistance, low process temperature and easy processing. Therefore, it would be advantageous to fabricate E-M devices from polymeric materials. Unfortunately, traditional piezoelectric polymers exhibit a much lower E-M performance than crystals and ceramics. It therefore makes sense to improve the E-M performance of the polymeric materials. In last decade, there have been breakthroughs that have led to major improvements in the E-M performance of polymeric materials. The newly developed E-M polymers, which are also called electroactive polymers (EAPs) or artificial muscles, exhibit a much better E-M performance than traditional E-M polymers and better than either crystals or

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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ceramics [1]. These discoveries open new avenues for the development of high performance E-M devices, such as actuators, sensors and transducers. More importantly, the high E-M performance observed in some EAPs may originate from changes at the molecular level, which makes it possible to develop high performance, electroactive micro/nanoscale and molecular devices such as microelectromechanical systems/nanoelectromechanical systems (MEMS/NEMS). In this chapter, the fundamentals of E-M materials and the properties of typical E-M polymers are reviewed, after which some challenges and issues related to the application of these polymers are discussed.

16.2 Fundamentals of Electromechanical Materials 16.2.1

Piezoelectric Effect

‘Piezo’ is derived from the Greek piezein, which means to squeeze or press. The piezoelectric effect is the phenomena where some materials generate an electric potential in response to an applied mechanical stress. Piezoelectricity was discovered by Pierre and Jacques Curie in 1880. Based on crystal symmetry and Neumann’s principle, it is known that the piezoelectric effect only exists in crystals that do not have a center of symmetry. There are 20 classes of crystal point symmetry groups that exhibit the piezoelectric effect. The piezoelectric effect can be described by [2]: x ¼ dE

D ¼ dX

ð16:1Þ

where x and X are the mechanical strain and stress, and E and D are electric field and displacement. Hence, the mechanical/electric response of a piezoelectric material to an electric/mechanical stimulus is linearly dependent on the stimulus. The constant, d, used in Equation (16.1) is the piezoelectric coefficient. The effect described in the first equation is often referred to as the converse piezoelectric effect, while that in the second is the direct piezoelectric effect. As E and D are first rank tensors and x and X are second rank tensors, d is a third rank tensor. The full tensor form of the piezoelectric constitutive equations can be written by adding the linear elastic (Hook’s law) and dielectric responses to Equation (16.1) [1]: xij ¼ dkij Ek þ sEijkl Xkl

Di ¼ eXik Ek þ dikl Xkl

ð16:2Þ

where sEijkl is the elastic compliance measured under constant E, eXij is the dielectric permittivity measured under constant X, and i, j, k, l ¼ 13. Actually, D and x can also be used as independent variables. By choosing one of the electric variables (E and D) and one of the elastic variables (X and x) as the two independent variables, different piezoelectric constitutive equations can be written to describe the piezoelectric response under different conditions. In addition to Equation (16.2), there are three more sets of piezoelectric constitutive equations. To simplify the calculation, matrix notation is usually used. In the matrix notation, ij or kl in tensor form is replaced by p or q according to: 11!1, 22!2, 33!3, 23 or 32 !4, 13 or 31 !5, 12 or 21 !6. Thus, Equation (16.2) transforms to:

Piezoelectric and Electrostrictive Polymer Actuators: Fundamentals

xp ¼ dkp Ek þ sEpq Xq

Di ¼ eXij Ej þ diq Xq

321

ð16:3Þ

Based on the symmetry of a material and Neumann’s principle, the number of independent elements for each material property can be reduced [2]. For instance, for an unstretched polymer, like poly(vinylidene fluoride) (PVDF) poled along its 3-direction, its structure belongs to the point group 1m and its properties are: 0

s11 Bs22 0 1 0 1 B 0 0 0 0 d15 0 e11 0 0 Bs13 @ A @ A spq ¼B dip ¼ 0 0 0 d15 0 0 eij ¼ 0 e11 0 B0 B d31 d31 d33 0 0 0 ; 0 0 e33 ; @0 0

s12 s22 s13 0 0 0

s13 s13 s33 0 0 0

0 0 0 s44 0 0

0 0 0 0 s44 0

1 0 0C C 0C C 0C C 0A s66

where s66 ¼ 2(s11 – s12). For a polymer stretched along its 1-direction and poled along its 3-direction, its structure belongs to the point group mm2 and its properties are: 0

s11 Bs12 0 1 0 1 B 0 0 0 0 d15 0 e11 0 0 Bs13 @ A @ A dip ¼ 0 0 0 d24 0 0 eij ¼ 0 e22 0 spq ¼B B0 B d31 d32 d33 0 0 0 ; 0 0 e33 ; @0 0

s12 s22 s23 0 0 0

s13 s23 s33 0 0 0

0 0 0 s44 0 0

0 0 0 0 s55 0

1 0 0C C 0C C 0C C 0A s66

It should be noted that the dielectric and elastic properties of a piezoelectric material are strongly dependent on its mechanical and electrical condition. For example: E X sD ijkl  sijkl ¼ dmij dnkl mn

exij  eXij ¼ dikl dimn cEklmn

ði; j; k; l; m; n ¼ 1  3Þ

ð16:4Þ

ði; j; k; l; m; n ¼ 1  3Þ

ð16:5Þ

where mn is defined as eik jk ¼ ij, and cklmn is defined as sijkl cklmn ¼ (ij)(mn), where ij ¼ 1 if i ¼ j and ij ¼ 0 if i 6¼ j.

16.2.2

Electrostrictive Effect

Unlike the piezoelectric effect, the electrostrictive effect occurs in all substances, whether crystalline or amorphous, solid or liquid. The electrostrictive strain response can be written as: xij ¼ Qijkl Pk Pl ¼ Mijkl Ek El

ði; j; k; l ¼ 1  3Þ

ð16:6Þ

where P is the electric polarization, Qijkl is the charge-related electrostrictive coefficient and Mijkl is the field-related electrostrictive coefficient. Both Qijkl and Mijkl are fourth rank

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Biomedical Applications of Electroactive Polymer Actuators

tensors. The value of Qijkl for a material is constant since Qijkl is an intrinsic material property. For a linear dielectric polymer: P ¼ ðe  e0 ÞE

ð16:7Þ

where e0 is the vacuum permittivity. Equation (16.7) links M and Q using the permittivity. However, most high performance E-M polymers are nonlinear dielectrics, so the permittivity defined by Equation (16.7) is dependent on the strength of the electric field. In this case, the value of M is dependent on the strength of the electric field. For an isotropic polymer under an electric field along the 3-direction (i.e. E ¼ E3) with a polarization response P along the 3-direction (i.e. P ¼ P3): x33 ¼ Q3333 P2 ¼ M3333 E2

and

x11 ¼ Q1133 P2 ¼ M1133 E2

ð16:8Þ

x3 ¼ Q33 P2 ¼ M33 E2

and

x1 ¼ Q13 P2 ¼ M13 E2

ð16:9Þ

or:

where x3 is the strain along the P direction, known as the longitudinal strain, and x1 is the strain perpendicular to the P direction, known as the transverse strain. This is different from the linear relationship between the field-induced strain and the electric field observed in piezoelectric polymers. The relationship in the electrostrictive material is quadratic, as shown in Figure 16.1. In most substances, the electrostrictive strain (M < 10–20 m2/V2) is too small to be used in practical applications. The first useful electrostrictive strain (~0.1 %) was found in relaxor ferroelectric ceramics [3], while newly developed electrostrictive polymers exhibit an electrostrictive strain of more than 5 % [1]. It is experimentally and theoretically proven

Electrostriction

x ∆x

ED

E

Piezoelectric

P

PD

∆P E

Figure 16.1 Schematic view of strain response in piezoelectric and electrostrictive polymers. When a small AC electric field superimposed on a DC bias field is applied on an electrostrictive polymer, an apparent piezoelectric response is observed.

Piezoelectric and Electrostrictive Polymer Actuators: Fundamentals

323

that Q33 (M33) <0 and Q13 (M13) >0 for an isotropic polymer and that Q33 (M33) >0 and Q13 (M13) <0 for an isotropic crystal or ceramic [4, 5]. Hence, an increase in E or P will result in a contraction of polymers, but an extension of crystals/ceramics, along the polarization direction. Based on thermodynamics, the value of Q reflects the change in the dielectric permittivity caused by mechanical stress [6]. The relation between Q and e is: Qijkl ¼ 

1 @ ð1=eij Þ 2 @Xkl

ð16:10Þ

Thus, the electrostrictive effect originates from the change in the dipole moment or density in a material. When a small AC electric field is imposed on a DC electric bias field, an apparent piezoelectric effect can be obtained in an electrostrictive polymer, as shown in Figure 16.1. For the field along the 3-direction, the dominating AC strain response term (Dx) is obtained using Equation (16.9) as: Dx ¼ 2Qp3 PD DP

ðp ¼ 6Þ

ð16:11Þ

where PD is the P induced by the DC bias field and DP is the polarization change induced by the AC field. Under a weak AC field, DP ¼ (ee0) DE, which leads to: Dx ¼ 2Qp3 ðe  e0 ÞDE

ð16:12Þ

Equation (16.12) is valid for both linear and nonlinear dielectric polymers as long as DE is small. Considering the similarity between Equations (16.1) and (16.12), an effective piezoelectric coefficient is introduced for electrostrictive polymers under a DC bias: d ¼ 2QPD ðe  e0 Þ

ð16:13Þ

For an isotropic polymer with a PD along the 3-direction, Equation (16.13) leads to: d33 ¼ 2Q33 PD ðe  e0 Þ

and

d31 ¼ 2Q13 PD ðe  e0 Þ

ð16:14Þ

where d33 and d31 are the piezoelectric coefficients along and perpendicular to the induced polarization, respectively. 16.2.3

Other Effects

When a dielectric material is under an electric field, positive and negative charges appear on its surfaces. The Coulombic force between these charges generates a stress on the material that results in a strain response, as shown in Figure 16.2. This stress is the Maxwell stress and it induces a strain approximately proportional to the square of the field. This is considered another E-M mechanism [7, 8].

324

Biomedical Applications of Electroactive Polymer Actuators Electric Field On

Electric Field Off V

Polymer film

Electrodes on top and bottom surfaces

Figure 16.2 effect.

Illustration of the electric-field-induced strain in soft polymer due to the Maxwell

For an isotropic polymer, the thickness (x3) and traverse (x1) strain responses due to the Maxwell effect are: x3 ¼ 

e 1 þ 2 2 E Y 2

x1 ¼

e 2 E 2Y

ðx  1Þ

ð16:15Þ

where Y ( ¼ 1/s11) and n ( ¼ s12/s11) are the Young’s modulus and Poisson ratio of the material, respectively. For most materials, the strain response generated by the Maxwell effect is too small to be useful in practical devices due to the high Young’s modulus. However, for soft polymers the Maxwell effect can generate useful strain and even giant strain response. For example, a field-induced strain of more than 100 % has been observed in some elastomers [8]. For a large strain, the relationship between the thickness (L) and the applied voltage (V) is:  2   L0 V Yln ð16:16Þ ¼e L L where L0 is the initial thickness. Based on Equations (16.9) and (16.15), the Maxwell effect and electrostrictive effect result in the same relationship between the strain and electric field and they therefore share some common features. For instance, an apparent piezoelectric effect can be observed when a DC bias is applied; the strain response can be enhanced by the nonuniformity of the electric field, which can be created either by employing nonuniform materials or by the presence of the space (trapping) charge. Due to the electrostrictive effect and the appearance of the space charge, an insulation material can exhibit piezoelectricity and is known as an electret [9, 10]. The piezoelectric constant of an electret depends on the space charge and its distribution as well as the nonuniformity in the elastic properties and electrostrictive coefficient of the materials.

16.3 Material Properties Related to Electromechanical Applications For an E-M device, there are various considerations regarding the material properties. Besides their piezoelectric constants and electrostrictive coefficients, many other properties are also critical to the E-M performance.

Piezoelectric and Electrostrictive Polymer Actuators: Fundamentals

16.3.1

325

Electromechanical Coupling Factor (k)

The electromechanical coupling factor (k) measures the ability of an E-M material to convert energy between the electrical and mechanical forms and is defined as [2]: k2 ¼

converted mechanical energy converted mechanical energy ¼ input electric energy input electric energy

Depending on the direction of the electric field and the directions associated with strain/ stress, there are many coupling factors. For a piezoelectric effect with electric field along direction i and the strain associated direction p, the k is: k2ip ¼

d2ip eXii sEpp

ð16:17Þ

If the electric field is applied along the 3-direction, the coupling factor is k33 when the actuation is along the 3-direction and k31 when the

actuation is along the 1-direction, which are k233 ¼ d233 = eX33 sE33 and k231 ¼ d 231 = eX33 sE11 , respectively. For the electrostrictive effect, the k is derived based on the induced polarization level P and strain x induced by a given electrical field [11]: k2ip ¼ sD pp

"

g  x2p    2 !# P S þ PE PE PE ln þ PS ln 1  P S  PE PS

ð16:18Þ

where i ¼ 1 or 3, corresponding to the transverse or longitudinal direction, and xp and PE are the strain and polarization responses, respectively, for the material under an electric field E. In Equation (16.18), it is assumed that the polarization-field (P-E) relationship follows |PE| ¼ PS tanh(g|E|), where PS is the saturation polarization and g is a constant. Therefore, for an electrostrictive polymer, the k is dependent on the applied electric field. Although k2 measures the capability of an E-M material to convert energy, it is not the energy conversion efficiency since the unconverted input energy is still stored in the material. Additionally, the value of k2 is not related to heat generation. The heat generation in an E-M material is determined by its dielectric loss and elastic loss. For piezoelectric ceramics, the mechanical loss is much smaller than the dielectric loss. Thus, the heat generated in piezoelectric ceramics is mainly determined by the dielectric loss. However, for E-M polymers, especially newly developed electrostrictive polymers, the elastic loss can be significantly higher than the dielectric loss. The k is also related to the coefficients measured under different conditions. For instance, the elastic compliance sE33 and sD constant field and charge, 33 (measured
under 2 E respectively) are related to each other as sD ¼ 1  k . Thus, under different external s 33 33 33 electric boundary conditions, a polymer with a large k will see a large difference in the elastic compliance. This can be used to tune the elastic modulus of the polymeric material by varying the electric conditions. For example, the elastic compliance of the electrostrictive material decreases with the DC bias. Similarly, ex33 ¼ ð1  k233 ÞeX33 , which means the dielectric permittivity measured under constant strain is smaller than that under constant stress.

326

16.3.2

Biomedical Applications of Electroactive Polymer Actuators

Elastic Response

For electromechanical and actuation applications, the following parameters are critical: maximum electric-induced strain response (xm), block force (Fb) and elastic energy density, including the volumetric elastic energy density (Wv) and gravimetric elastic energy density (Wg). The xm of an E-M material is directly related to the displacement generated in an actuator. For most piezoelectric ceramics and polymers, xm is about 0.1–0.2 %, while the newly developed E-M polymers exhibit a strain response of more than 5%, in some cases achieving as much as 100 %. This makes it possible to create actuators that exhibit a giant displacement. Fb measures the maximum force needed to maintain the zero displacement when the material is under electric field. For an E-M polymer with a linear elastic response, Fb can be expressed as Fb ¼ Yxm. Due to their low Young’s modulus, E-M polymers usually exhibit a small block force compared to E-M ceramics. The elastic energy density characterize the elastic energy stored in the E-M materials and Yx2 Yx2 are defined as Wv ¼ 2 m , and Wg ¼ 2m , where r is the density of the material. Wv is related to the volume of the device, while Wg is related to the mass of the device. A device made of an E-M material with a higher elastic energy density would have a smaller size/mass. As shown in Tables 16.1 and 16.2, the newly developed electrostrictive polymers exhibit a much higher elastic energy density than piezoelectric ceramics and polymers. 16.3.3

Frequency and Temperature Responses

The frequency and temperature dependence of E-M performance in polymers is much stronger than that in ceramics at frequencies lower than 107 Hz. For the E-M polymers, the dielectric constant decreases with frequency due to dielectric relaxation. The relaxation frequency of the E-M polymers is in the kHz to MHz range, depending on the materials. At frequencies around the relaxation frequency, there is a high dielectric loss. Additionally, as most E-M polymers exhibit a strong elastic relaxation that results in a higher Young’s modulus at a higher frequency, all the E-M properties are frequency dependent. For instance, the xm decreases with increasing frequency. Besides the frequency dependence of the properties of these materials, the resonance frequency of an E-M device such as an actuator plays an important role in the frequency dependence of its performance. At the resonance frequency, xm can be much higher than at a low frequency. The temperature has a very complicated influence on the E-M performance of polymers. The Young’s modulus of polymer decreases with the temperature, but the temperature dependence of its dielectric properties can be very different. For some of the piezoelectric and electrostrictive polymers, there is a phase transition at temperatures close to room temperature. At the phase transition temperature, the dielectric permittivity reaches its maximum and the polymer exhibits a high dielectric loss. Additionally, at temperatures around the glass transition temperature, all polymers exhibit some dielectric relaxation and elastic relaxation, resulting in a frequency dependence of material properties and a high dielectric and elastic loss. Due to the dielectric and elastic relaxations, the E-M response in polymers often shows strong relaxation behavior. The coefficients in the piezoelectric constitutive

Piezoelectric and Electrostrictive Polymer Actuators: Fundamentals

327

Table 16.1 The thickness strain response and corresponding electromechanical performance of different materials at room temperature Materials

Y (GPa)

xm (%)

E (V/mm)

Wv (J/cc)

Fb (MPa)

k33

0.13

128

0.6

Ceramics [12]

64

0.2 ~100

Polymers [12]

3

0.1 ~ 40

0.0015 3

0.2

Irradiated P(VDFTrFE) [18, 19, 24]

0.4

5

150

0.5

20

0.3

P(VDFTrFECTFE) [25, 26]

0.4

4

150

0.32

16

0.28

62/38/4

1.2

4.5 130

1.22

54

0.55

68/32/9

0.3

7

170

0.74

21

P(VDFCTFE) [28, 29]

0.82

5.5 215

1.24

45.1

Grafted elastomer [30, 31]

0.55

4

120

0.44

22

Polyurethane [32, 33]

0.017

11

100

0.1

1.9

0.01

0.03

P(VDFTrFECFE) [26, 27]

0.0036

7.5 11

Remarks

Piezoelectric materials

Electrostrictive polymers

0.39

Maxwell stress based polymers

Acrylic (VHB 4910) [8, 34]

0.001

68

239

~0.1

0.7

0.79

Silicone (CF19–2186) [8, 34]

0.001

39

181

0.2

0.6

0.63

Table 16.2 The transverse strain response and corresponding electromechanical performance of different materials at room temperature Materials

Y (GPa)

xm (%)

E (V/mm)

Ceramics

64

–0.1

~100

Polymers

3

0.1

~40

Irradiated P(VDF-TrFE) [19, 24]

1.1

4.5

Acrylic (VHB 4910) [8, 34]

0.001

Silicone (CF19–2186) [8, 34]

0.001

Conventional piezoelectrics [12]

Wv (J/cc)

Fb (MPa)

k31

0.03

128

0.3

0.0015

3

0.2

85

1.1

49.5

0.65

215

239

2.4

2.2

0.85

63

181

0.2

0.6

0.63

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Biomedical Applications of Electroactive Polymer Actuators

equation are frequency dependent and complex as shown in Equation (16.19): d ip ¼ d0ip  jd 00 ip ;

sip ¼ s0ip  js00 ip ;

and

eij ¼ e0ij  e00 ij ;

ð16:19Þ

where j represents an imaginary unit.

16.4 Typical Electromechanical Polymers and Their Properties 16.4.1

Piezoelectric Polymers

Polarization (mC/m2)

Ferroelectrics are an important group of piezoelectric materials because they exhibit a much higher piezoelectric response than nonferroelectric piezoelectrics. For instance, for ceramics, d is 100–101 pC/N for nonferroelectric piezoelectrics and 102–103 pC/N for ferroelectrics; for polymers, d is 10–1–100 pC/N for nonferroelectric piezoelectrics and 100–101 pC/N for ferroelectrics. This ferroelectricity is defined by the appearance of a switchable spontaneous polarization and characterized by a characteristic P-E loop, as shown in Figure 16.3. Besides the spontaneous polarization, the remnant polarization (Pr) and coercive field (Ec) are also critical parameters for a ferroelectric. Compared to ceramics, ferroelectric polymers exhibit a smaller Pr and a much higher Ec.

100

0

–100 –200

Pr

Ec

Ec Pr

–100 0 100 Electric Field (V/µm)

200

Figure 16.3 Polarization versus electric field (P–E) loop for a ferroelectric polymer – P(VDF– TrFE), where Pr and Ec are the remnant polarization and coercive field.

Ferroelectricity has been observed in many polymers [12], such as PVDF, copolymers of VDF with trifluoroethylene (TrFE) or tetrafluoroethylene (TFE) and odd-numbered polyamides (nylons). It occurs mostly in the form of semicrystalline polymers, in which the crystallites are embedded in an amorphous surrounding. To be ferroelectric, the polymer must have the following characteristics: (i) the repeating unit of the polymer chain should be strongly polar so that a large permanent dipole moment is associated with the unit, for example, in PVDF the repeating unit is CH2–CF2, which has a dipole moment of about 1 Debye, and in nylon, the repeating unit has both N–H and C ¼ O polar groups; (ii) the conformation of the polymer chain enhances the dipole moment of repeating units by aligning dipoles along the same direction. This explains why even-numbered nylon is not ferroelectric but odd-numbered nylon is; and (iii) the polymer chains are packed into a

Piezoelectric and Electrostrictive Polymer Actuators: Fundamentals

329

crystal structure in which the dipole moments of different polymer chains are aligned. For example, PVDF in -phase is ferroelectric while the -phase is non-ferroelectric [13], as shown in Figure 16.4, where the conformation of the polymer chain and the crystal structure of PVDF are illustrated.

Figure 16.4 Illustration of the influence of the polymer chain conformation and the crystalline structure on the dipole moment and ferroelectricity of PVDF: (left) polymer chain takes all-trans conformation, resulting in a net dipole moment for the polymer chain. These polymer chains ˚ , b ¼ 4.91 A ˚ and c ¼ 2.58 A ˚ along are packed into a ferroelectric structure (b phase: a ¼ 8.58 A chain direction) in which the dipole moments line up; (right) polymer chain takes trans-gauche conformation, resulting in a zero net dipole moment. These polymer chains are packed into a ˚ , b ¼ 9.64 A ˚ , and c ¼ 4.62 A ˚ along the chain non-ferroelectric structure ( phase: a ¼ 4.96 A direction) in which the dipole moments cancel each other out.

Due to the appearance of ferroelectric domains, ferroelectric polymers have to be poled to obtain piezoelectricity. The poling conditions, including the electric field and temperature, is critical to the resulting piezoelectricity of the polymer. Different methods are used to pole the polymers. Two widely used methods are electrode poling, in which the poling electric field is applied through two metal electrodes, and corona poling, in which a high electric field is applied directly to the polymer film without metal electrodes. The former is easy to conduct in a laboratory, but the latter is more efficient because if localized electric breakdown occurs, this will not affect the corona poling process. Since the ferroelectricity in these polymers originates from the crystallites, a higher crystallinity usually results in a higher piezoelectric response in these polymers. In order to achieve a higher crystallinity and a better orientation effect in the ferroelectric polymers, mechanical stretching, including onedimensional and two-dimensional stretching, is widely employed. The mechanical stretching is usually conducted prior to the thermal annealing process that increases the crystallinity of the polymer. Sometimes, a stress along the stretching direction is employed during thermal annealing to maintain the stretching effect. In these semicrystalline polymers, the interfacial layer between the crystallites and amorphous regions has a strong influence on the poling process and the E-M performance.

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Among all ferroelectric polymers, PVDF and P(VDF–TrFE) are the best known and exhibit the highest piezoelectric response (d33 ~–35 pC/N and d31~28 pC/N, k33~0.15, k31~0.15, e33/e ffi 13, Y ¼ 2.5 GPa, n ¼ 0.39 and r ¼ 1.78–1.80 g/cm3) at room temperature, while the odd-numbered nylons exhibit a higher piezoelectric response (d31 ¼ 15~20 pC/N, k31 ~0.05) at high temperatures (up to 200 °C) due to their high glass transition temperatures (60~190 °C). It should be mentioned that the nylons exhibit a particularly high dielectric loss at high temperatures. In an effort to improve the piezoelectric response in P(VDF–TrFE), a single crystal polymer film of P(VDF–TrFE) with 75 mol% of VDF was studied and the film found to have the best piezoelectric performance of piezoelectric polymers to date (d33 ¼ –30 pC/N, k33 ffi 0.33) [14, 15]. Polymer blends have also been explored to improve the piezoelectric performance and thermal stability of these polymers. Although the physical mechanism is different, electrets exhibit an apparent piezoelectric response. The piezoelectric constants of some electrets can be comparable to those of ceramics. Currently, polymer electrets are widely employed in microphones and other E-M devices due to their low cost [16, 17]. To improve the piezoelectric response in polymer electrets, air was recently introduced as one of the constituent materials, which results in porous polymers. A very high piezoelectric response (d33 up to 1000 pC/N, Y ¼ 1~10 MPa) was reported in these porous polymers [9, 10]. Here, each air bubble can be charged by discharging air, which makes air bubble a dipole with a very large dipole moment. Interestingly, these porous electrets can exhibit an apparent ferroelectricity due to the charging and discharging activity in the air bubbles in the porous polymer and are therefore referred to as ferroelectrets [9, 10]. 16.4.2

Electrostrictive Polymers

Many electrostrictive polymers have been developed in the last decade [1] and these newly developed electrostrictive polymers exhibit a high electric-field-induced strain, as shown in Figure 16.5a, where the maximum thickness strain response of the polymer at different fields is given. A typical relationship between the strain response and the electric field observed in these polymers is shown in Figure 16.5b. All these electrostrictive polymers are polar polymers that contain polar units in the polymer chain. The electrostrictive strain response reflects the change in these polar units due to an electric field. The first reported electrostrictive polymer with a high strain response was the highenergy-electron irradiated P(VDF–TrFE) copolymer [18, 19]. The irradiation converts the polymer from a normal piezoelectric to an electrostrictive state. Based on the experimental results, the Q33 of the irradiated P(VDF–TrFE) can range from –4 to –15 m2/C (the M is greater than 2  10–18 m2/V2), depending on the sample processing conditions. Additionally, irradiated copolymers exhibit a high dielectric permittivity (e/e0 ~60) at room temperature. It has also been reported that the irradiated copolymers are actually relaxor ferroelectrics. Similar results were observed in high-energy proton irradiated P(VDF–TrFE) (M33 ¼ –1.83  10–18 m2/V2). By varying the film processing conditions, the transverse strain in an irradiated P(VDF– TrFE) copolymer can be tuned over a wide range. For instance, the transverse strain response in unstretched films is relatively small (~ þ1 % at ~100 MV/m), with an

Piezoelectric and Electrostrictive Polymer Actuators: Fundamentals 0

Irradiated P(VDF-TrFE)

–4 –2 P(VDF-CTFE)

0

Thickness Strain (%)

Thickness Strain (%)

P(VDF-TrFE-CFE)

–6

331

–1 –2 –3 –4 –5

0

100 200 Electric Field (V/µm) (a)

–100 0 100 Electric Field (V/µm) (b)

Figure 16.5 (a) The maximum thickness strain at different electric fields for three newly developed electrostrictive polymers. (b) The relationship between the strain response and electric field for a typical electrostrictive polymer

amplitude ratio of the transverse to longitudinal strain of less than 0.33, while the transverse strain in stretched films can reach þ4.5 % along the stretching direction. For the stretched film, a high Young’s modulus is also obtained along the stretching direction, which results in a high k (k33 > 0.3, k31 > 0.65) in the irradiated and stretched films. The k31 observed in the irradiated copolymer is higher than that in ceramics. Microstructure studies reveal that the high electrostrictive response observed in the irradiated copolymer originates from the electric-field-induced conformation change. The defects induced by the irradiation result in nano-sized crystallites with a highly stressed interfacial layer, which make the trans-gauche conformation favorable. In contrast, the application of an electric field favors the all-trans conformation. Consequently, an external electric field can be used to switch the gauche conformation to the trans conformation, producing a large strain response. For the lattice shown in Figure 16.4, a strain of 9% along the polymer chain is imposed by the gauche-to-trans conformation switch. Inspired by the irradiated P(VDF–TrFE) copolymers, different approaches, especially a chemical approach, have been investigated to introduce defects into PVDF based polymers. For example, chlorofluoroethylene (CFE), chlorotrifluoroethylene (CTFE) and hexafluoropropylene (HFP) have been introduced as bulky monomers into P(VDF– TrFE) to form terpolymers. These bulkier monomers make the polymer chain favor the gauche conformation. As expected, under a high electric field these terpolymers exhibit a high electrostrictive strain response. Among these terpolymers, P(VDF–TrFE–CFE) terpolymers exhibit the best E-M performance. The relatively low E-M performance observed in P(VDF–TrFE–HFP) is probably due to its low crystallinity since the HFP group is too big to cocrystallize with VDF and TrFE. A detailed study of P(VDF–TrFE–CTFE) terpolymer found that the conformation change of polymer chains in the crystallites was not sufficient to explain the experimentally observed high electrostrictive strain response [20]. It is believed that a partially ordered interfacial layer in the semicrystalline polymer plays an important role in the electrostrictive strain response [21]. The contribution of the partially ordered region to the electrostrictive strain response in these electrostrictive polymers was further confirmed by a study of P(VDF–CTFE) copolymers [22].

332

16.4.3

Biomedical Applications of Electroactive Polymer Actuators

Maxwell Stress Effect Based Polymers

A significant breakthrough in using the Maxwell stress effect for E-M polymers is the discovery that some soft polymers, such as dielectric elastomers (Y ~1 MPa), can exhibit a very high electric-field-induced strain response (>100 %) [8]. Interestingly, this high strain response observed in the elastomer was under a high pre-strain (~300 %). Experiments have revealed that the high pre-strain increases the breakdown field that makes it possible to apply a very high field onto the elastomer (>200 MV/m) to generate a high strain response. A detailed study revealed that the strain response of the elastomer under the same field increases initially with the pre-strain and then decreases [23], indicating there is an optimum pre-strain that allows the elastomer to exhibit a giant strain response. It is thought that the pre-strain creates partially ordered regions in the elastomer and that these partially ordered regions respond to the external electric field due to the dipole associated with the polymer chain. 16.4.4

Practical Considerations

The electromechanical properties related to actuator applications of newly developed electroactive polymers are summarized in Tables 16.1 and 16.2. For comparison, the piezoelectric ceramics and polymers are also included. Clearly, the newly developed electroactive polymers exhibit a much higher energy density than the traditional piezoelectrics. However, due to their low Young’s modulus values, the E-M polymers exhibit a smaller block force. It is interesting to note that the coupling factor of some polymers is comparable or even higher than that of piezoelectric ceramics. It should be pointed out, however, that these polymers have a working frequency range that is usually smaller than for ceramics. As discussed above, dielectric relaxation and elastic relaxation are the key factors that limit the working frequency range. As indicated in Tables 16.1 and 16.2, a high electric field is usually required in order to achieve a high strain response in the newly developed E-M polymers. Although this limits the application of these polymers, it is very interesting to be able to reduce the electric field used to drive these E-M polymers.

16.5 Conclusions The unique properties of polymeric materials, namely their flexibility, light weight and easy processing, open up new possibilities for E-M devices. Recent developments have shown that E-M polymers can exhibit far better E-M performance than traditional polymers, crystals and ceramics [1]. However, there are also problems due to issues such as their low block force, limited range of frequencies and the requirement for a high electric field. Nevertheless, E-M polymers offer exciting new challenges for scientists and researchers.

References 1. Cheng, Z.-Y. and Zhang, Q.M. (2008) Field-activated electroactive polymers, MRS Bulletin, 33, 183–7. 2. Nye, J.F., Physical Properties of Crystals, Clarendon Press, Oxford, 1987.

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3. Cross, L.E., Jang, S.J., Newnham, R.E., et al. (1980) Large electrostrictive effects in relaxor ferroelectrics, Ferroelectrics, 23, 187–91. 4. Shkel, Y.M. and Klingenberg, D.J. (1998) Electrostriction of polarizable materials: Comparison of models with experimental data, J. App. Phys, 83, 415–24. 5. Xu, Y.H. Ferroelectric Materials and Their Applications, Elsevier Science, The Netherlands (1991). 6. Uchino, K., Nomura, S., Vedam, K., et al. (1984) Pressure dependence of the refractive index and dielectric constant in a fluoroperovskite, KMgF3, Phys Rev., B29, 6921–5. 7. Landau, L.D. and Lifshitz, E.M., Electrodynamics of Continuous Media, Pergamon Press, Oxford, 1970. 8. Pelrine, R., Kornbluh, R., Pei, Q. and Joseph, J. (2000) High-speed electrically actuated elastomers with strain greater than 100%, Science, 287, 836–9. 9. Bauer, S., Gerhard-Multhaupt, R. and Sessler, G. (2004) Ferroelectrets: Soft electroactive foams for transducers, Physics Today, 57, 37–44. 10. Wirges, W., Wagener, M., Voronina, O., et al. (2007) Optimized preparation of elastically soft, highly piezoelectric, cellular ferroelectrets from nonvoided poly(ethylene terephthalate) films, Adv. Funct. Mat., 17, 324–9. 11. Hom, C., Pilgrim, S., Shankar, N., et al. (1994) Calculation of quasi-static electromechanical coupling coefficients for electrostrictive ceramic materials, IEEE Transactions on Ultrasonics, Ferroelectrics and Frequency Control, 41, 542–51. 12. Nalwa, H.S. (1995) Ferroelectric Polymers: Chemistry, Physics, and Applications, Marcel Dekker, Inc, New York. 13. Lovinger, A.J. (1982) in Developments in Crystalline Polymers (ed.D.C. Bassett,), Applied Science Publishers, London, 195. 14. Omote, K., Ohigashi, H. and Koga, K. (1997) Temperature dependence of elastic, dielectric and piezoelectric properties of ‘single crystalline’ films of vinylidene fluoride trifluoroethylene copolymer, J. App. Phys, 81, 2760–9. 15. Ohigashi, H., Omote, K. and Gomyo, T. (1995) Formation of ‘single crystalline films’ of ferroelectric copolymers of vinylidene fluoride and trifluoroethylene, App. Phys Lett., 66, 3281–3. 16. Sessler, G.M. (1963) Electrostatic microphones with electret foils, J. Acoust. Soc. Am. 35, 1354–7. 17. Sessler, G.M. and West, J.E. (1973) Electret transducers: a Review, J. Acoust. Soc. Am. 53, 1589–1600. 18. Zhang, Q.M., Bharti, V. and Zhao, X. (1998) Giant electrostriction and relaxor ferroelectric behavior in electron irradiated P(VDF-TrFE), Science, 280, 2101–4. 19. Cheng, Z.-Y., Xu, T.-B., Bharti, V., et al. (1999) Transverse strain responses in the electrostrictive poly(vinylidene fluoride-trifluorethylene) copolymer, App. Phys Lett. 74, 1901–3. 20. Garret, J.T., Roland, C.M., Petchsuk, A. and Chung, T.C. (2003) Electrostrictive behavior of poly(vinylidene fluoride-trifluoroethylene-chlorotrifluoroethylene), App. Phys Lett. 83, 1190–2. 21. Li, Z.M., Li, S.Q. and Cheng, Z.-Y. (2005) Crystalline structure and transition behavior of recrystallized-irradiated P(VDF-TrFE) 65/35 copolymers, J. App. Phys, 97, 014102. 22. Li, Z.M., Wang, Y.H. and Cheng, Z.-Y. (2006) Electromechanical properties of poly(vinylidenefluoride-chlorotrifluoroethylene) copolymer, App. Phys Lett. 88, 062904. 23. Yang, G., Yao, G., Ren, W., et al. (2005) The strain response of silicone dielectric elastomer actuators, in Smart Structures and Materials 2005: Electroactive polymer actuators and devices (EAPAD) (ed.Y. Bar-Cohen,), Proceedings of SPIE, 5759, 134–43. 24. Cheng, Z.-Y., Bharti, V., Xu, T.B., et al. (2001) Electrostrictive poly(vinylidene fluoridetrifluoroethylene) copolymers, Sensors and Actuators A90, 138–47. 25. Xu, H., Cheng, Z.-Y., Olson, D., et al. (2001) Ferroelectric and electromechanical properties of P(VDF-TrFE-CTFE) terpolymer, App. Phys Lett.78, 2360–2.

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26. Huang, C., Klein, R., Xia, F., et al. (2004) P(VDF-TrFE-CFE) based high performance electroactive polymers, IEEE Transactions on Dielectrics and Electrical Insulation 11, 299–311. 27. Xia, F, Cheng, Z.-Y., Xu, H., et al. (2002) High electromechanical responses in a P(VDF-TrFECFE) terpolymer, Adv. Mat. 14, 1574–7. 28. Li, Z.M., Wang, Y.H. and Cheng, Z.-Y. (2006) Electromechanical properties of poly(vinyledene-fluoride-chlorotrifluoroethylene) copolymer, App. Phys Lett. 88, 062904. 29. Li, Z.M. (2004) Novel electroactive poly(vinylidene fluoride)-based polymer systems and their applications. PhD Thesis, Auburn University. 30. Su, J., Hales, K. and Xu, T.B. (2003) Composition and annealing effects on the response of electrostrictive graft elastomers, Proceedings of SPIE, 5051, 191–7. 31. Su, J., Harrison, J.S. and St. Clair, T.L. (2005) Electrostrictive graft elastomers, US Patent 6,515,077. 32. Zhenyi, M., Scheinbein, J.I., Lee, J.W. and Newman, B.A. (1994) High field electrostrictive response of polymers, J. Polym. Sci. B: Polym. Phys 32, 2721–31. 33. Pelrine, R.E., Kornbluh, R.D. and Joseph, J.P. (1998) Electrostriction of polymer dielectrics with compliant electrodes as a means of actuation, Sensors and Actuators A44, 77-85. 34. Kornbluh, R., Pelrine, R., Pei, Q.B. et al. (2000) Ultrahigh strain response of field-actuated elastomeric polymers, Proceedings of SPIE, 3987, 51–64.

17 Miniature High Frequency Focused Ultrasonic Transducers for Minimally Invasive Imaging Procedures Aaron Fleischman, Sushma Srivanas, Chaitanya Chandrana and Shuvo Roy Cleveland Clinic, Lerner Research Institute, Department of Biomedical Engineering, Cleveland, USA

17.1 Introduction High frequency ultrasound has many medical applications and has been demonstrated to provide high quality imaging of the anterior chamber of the eye, skin imaging, intravascular imaging and endoluminal imaging [1, 2]. Image resolution is primarily determined by properties of the transducer such as bandwidth, focusing and pulse length. High resolution images can be obtained by using broad bandwidth materials and focusing the ultrasonic transducer. Though higher frequency transducers produce higher resolution images, it is at the cost of reduced penetration [3]. This has limited high frequency ultrasonic transducers to areas of interest that require small amounts of penetration such as Intravascular Ultrasound (IVUS). However, IVUS and endoluminal imaging applications in general have not benefited from the increased lateral resolution that can be obtained by the focusing of the transducer because of the difficulties associated with the fabrication of small lenses, mirrors or shaped transducers that can fit within a catheter profile.

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Figure 17.1 Illustration of rotating IVUS geometry used to generate intravascular images.

A schematic diagram of a typical rotating IVUS probe is shown in Figure 17.1. The transducer comprises an active element with a diameter of about 0.6 mm located on a rotating core contained within a plastic sheath. An image is formed by mechanically rotating the transducer in a complete turn around the interior of the artery while repeating the pulse echo sequence at each 1.2 step for a total of 256 scan lines per rotation. The catheter is pulled back at a controlled rate to scan the desired portion of the coronary artery. Image processing has been used to assemble a ‘fly through’ depiction of the artery. Overall image quality is fundamentally limited by inherent characteristics of lead zirconate titanate (Pb[ZrxTi1 x]O3 0<x<1) (PZT) transducer designs. IVUS imaging is a technology that permits tomographic visualization of a cross-section through the vessel wall and can assist physicians in selecting and evaluating types of therapeutic interventions [4, 5]. Recent clinical studies have documented its sensitivity in detecting atherosclerosis, and it is increasingly employed to assist in finding the most appropriate therapeutic intervention [4–10]. IVUS has been used to detect proper stent apposition, late stent thrombosis and plaque prolapse [11]. Several studies have compared geometric parameters (i.e. areas and thickness of gross features) measured by histological evaluation or angiography vs IVUS [12–14]. These studies have dispelled much of the initial skepticism toward the gross accuracy and reliability of IVUS images and caused clinical experts to take a more critical look at the limitations of the traditional imaging modality of angiography. Comparisons of IVUS results with those from histology resulted in regression lines closer to unity [15–17]. The development of 3-D segmentation techniques, taking advantage of the similarity of borders between frames, has been a recent evolution in the movement toward routine quantitative analysis in catheterization laboratories [18–24]. Correct identification of plaque types depends on their structure, composition and sufficient image resolution. Recent work in IVUS backscatter analysis demonstrates the capability of IVUS to characterize specific lesions and identify the kinds of plaques that lead to various clinical syndromes [25, 26]. Other techniques, such as wavelet analysis and elastography [27], are being developed to identify plaque composition as well. These techniques have been used to identify plaque composition and geometry using IVUS. The advances in IVUS have allowed for automated quantitative analysis of vessel

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geometry and plaque compositions. These advances have made IVUS more practical to use by enabling reproducible measurements and characterization of coronary lesions in virtually real time. In fact, IVUS has become the standard by which to judge the efficacy of many pharmacologic interventions and the understanding of artherosclerosis [28]. This has resulted in the increased use of this procedure by cardiologists and has grown from 10 to 15 % of cases requiring percutaneous coronary intervention (PCI) [29]. This has primarily been driven by physicians’ realization that the incidence of late stent thrombosis may be related to stent malapposition, underexpansion or ‘geographic miss’ (failure to fully cover a diseased arterial segment with stent) [30].

17.2 Coronary Imaging Needs Recent advances in the understanding of cardiac syndromes and the therapies used to treat them require higher resolution images with which to diagnose various lesions and to evaluate therapeutic interventions. Diagnosing and understanding in areas such as the role that vulnerable plaques play in sudden myocardial infarctions (MIs), complications arising from poor stent deployment, late stent thrombosis associated with drug-eluting stents and the desire to evaluate neointimal coverage of stents have made increased resolution a true clinical need. The importance of these areas is clear but limitations in imaging have prevented studies that quantify these issues, making it difficult to attribute absolute minimum resolutions required to study these issues. But, qualitatively, the direction of coronary imaging is apparent – increased resolution will allow better images and geometric definition of the plaque components to evaluate the composition of a lesion and closer inspection of the stent–artery interface to evaluate stent apposition, as well as thrombosis and neointimal coverage. 17.2.1

Vulnerable Plaques

Coronary heart disease, acute coronary syndrome and angina pectoris are the leading cause of mortality in the United States. Annually, 451 000 people die from these diseases, 800 000 suffer acute MI and 156 000 died from an MI, and many more are hospitalized due to an acute manifestation of ischemic heart disease. The personal, social and economic consequences are enormous, for 2008 they are estimated at $156.4 billion annually in the United States alone [31]. Erosion and rupture of vulnerable atherosclerotic plaque is the cause of most acute coronary syndromes [32, 33]. Plaque rupture leads to the formation of an intracoronary thrombus, which produces an obstruction that acutely limits coronary artery blood flow, resulting in myocardial ischemia or necrosis. Multiple clinical and autopsy studies have confirmed the pathogenic role of coronary thrombus in cases of acute MI, unstable angina and sudden cardiac death [33, 34]. The lesions that harbor vulnerable plaques are often mildly stenotic on angiographic examination and, consequently, their stability cannot be assessed. The stability of atherosclerotic plaques is related to histological composition Figure 17.2. Unstable plaques typically comprise thin (<65 mm) fibrous caps infiltrated with macrophages that encapsulate lipid-rich necrotic cores with adjacent microcalcification [33, 34].

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Figure 17.2

Histological image of a vulnerable plaque.

Lesion geometry as a predictor of plaque rupture lies at the core of an enigma among investigators today, who tend to fall into two schools of thought: (i) post mortem studies have consistently shown that culprit lesions harbor large atherosclerotic plaques [35–38]; (ii) on the other hand, several investigators have studied patients with MI in whom an angiogram was performed within one year prior to a coronary event [39–41]. Each of these studies reported that prior angiography most frequently demonstrated a stenosis of less than 50 % at the culprit site responsible for subsequent occlusion. On the basis of these observations, some investigators have proposed that plaques that cause acute coronary syndromes are small, ‘early’ lesions. This assumption does not take into account that positive remodeling attenuates the encroachment of the plaque into the lumen, thereby maintaining the luminal area. [42–44]. A study by Tuzcu and colleagues has demonstrated an association between the direction of arterial remodeling and the type of clinical presentation in patients with coronary artery disease [45]. Positive remodeling of the culprit lesion was associated with unstable clinical presentation, whereas negative remodeling was more common in patients with stable clinical presentation. Further, lesions at higher risk have a thin and or fissured cap with increasing risk if it has a large lipid core and a history of rupture and remodeling [33]. It is not possible currently to identify coronary artery plaques that are vulnerable to rupture and stabilize them or otherwise prevent subsequent rupture. In fact, the current gold standard for the assessment of coronary artery disease, coronary angiography, is unable to identify those coronary artery atherosclerotic lesions that may produce future acute coronary syndromes; it can only identify the culprit lesions in symptomatic cases [12, 46–49]. In two thirds of patients presenting with acute ischemic syndromes, the culprit lesions were thought to be insignificant on a previous coronary angiogram [15]. Thus, the vulnerability of a lesion to rupture is thought to depend not on the degree of stenosis it produces within the coronary artery, but rather on its structure, composition and the

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mechanical forces acting upon it [32]. These factors cannot be ascertained by coronary angiography. This limitation of angiography is exemplified in that 33 % of acute MIs that occur following successful coronary artery bypass surgery are caused by occlusion of an artery that did not contain an angiographically significant stenosis at the time of surgery [16]. Similarly, 57 % of infarcts after percutaneous transluminal coronary artery angioplasty (PTCA) are caused by occlusion of an artery that did not contain a ‘significant’ stenosis at the time of PTCA [50]. In vivo identification of plaque composition with sufficient resolution should enable the detection of vulnerable plaque before rupture enabling some local treatment. Being able to image vulnerable plaques may lead to the detection of the characteristics of a vulnerable plaque that could allow the development of predictive models to assess the risk that a vulnerable plaque poses to the patient [33]. Current state-of-the-art IVUS systems, such as those that use VH from Volcano Corporation, do not measure the thin fibrous cap explicitly as the axial resolution is too poor to detect the thin caps. Rather, when a necrotic core is with in 125 microns of the lumen, it is assumed to be covered by a thin fibrous cap and is deemed vulnerable [29, 51]. 17.2.2

Stent Thrombosis

Stent thrombosis is the most devastating complication of drug-eluting stent (DES) implantation, occurring in prior studies in 0.4–3.5 % of patients, depending on baseline patient characteristics, definitions and duration of follow-up [52–62]. Following stent thrombosis, the incidence of death and MI is more than 80 %, with mortality rates in prior studies ranging from 20 to 45 % [52–55]; as such, the prevention of stent thrombosis is a major imperative. The clinical and angiographic risk factors for stent thrombosis have varied in different studies, with reduced left ventricular function, thrombus/acute coronary syndromes, bifurcation stenting, left main stenting, small vessels, long lesions/long stents, diabetes, renal insufficiency, prior vascular brachytherapy, suboptimal stent result and premature thienopyridine discontinuation having been most prominently emphasized [52–57]. In addition, recent analyses have suggested that implantation of a DES is associated with a higher rate of late stent thrombosis when compared with bare metal stents, and prolonged dual antiplatelet therapy (at least one year) is recommended after DES implantation in patients at low risk of bleeding. However, prolonged antiplatelet therapy is not a benign treatment strategy, and although it is meant to reduce bleeding risk in individuals compared to other regimens, treatment must still be discontinued prior to any surgical procedure. The decision to cease antiplatelet therapy can itself lead to thrombosis and death. Stent thrombosis and DES-related stent thrombosis is probably a multifactorial issue; contributing factors include stent malapposition, positive remodeling of the vessel and underexpansion of the stent during deployment [63–65]. Evaluating stent deployment, stent malapposition and late stent thrombosis requires imaging the stent in relation to the adjacent vessel wall. However, the poor lateral resolution of IVUS results in the bright signal from the stent struts ‘washing out’/obscuring the signal from the adjacent vessel wall [64]. Since ideally the tissue–stent boundary should be sharp, the better the lateral resolution the more useful the information to the clinician. It is generally believed that stent thrombosis is driven by incomplete endothelialization of the stent struts [66, 67]. IVUS lacks the required axial resolution to visualize neointimal

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formation around the stent, which may be as thin as 29 mm (1 Standard Deviation below the mean) with an average of 53 mm [66, 67]. Optical coherence tomography (OCT) has been able to detect neointimal stent coverage [66–68]. The development of a focused broad bandwidth ultrasound transducer with axial resolutions of 20 mm or higher, as proposed in this application, would be able to detect neointimal coverage. Having this capability would allow patients under consideration for antiplatelet therapy cessation to undergo IVUS imaging and determine the degree of neointimal coverage of the stent. Arterial thrombotic events are the leading cause of MI and have recently been identified as a major contributing factor to plaque progression. The detection of thrombus in a coronary artery is crucially important, as it is used to verify plaque rupture, evaluate the efficacy of thrombolytic therapy, assess the need for thrombectomy and determine the risk for distal embolization and thrombus protrusion through stent struts if percutaneous coronary intervention is warranted [69]. Unfortunately, no FDA-approved technology can accurately detect intraluminal thrombus.

17.3 High Resolution Ultrasonic Transducers The imaging performance of an ultrasonic transducer is related to the lateral resolution, the axial resolution and the depth of field of the transducer. The lateral resolution (Rlat) is the width of the ultrasonic beam in the focal region. The axial resolution (Rax) is determined by the pulse bandwidth. The depth of field (DOF) is the length of the focused region of the beam. These values can be approximated for a spherical radiator as [2]: Rlat ¼ lðf Rax ¼

numberÞ 1 C 2 BW

DOF ¼ 7:0 lðf

number Þ 2

ð17:1Þ ð17:2Þ

ð17:3Þ

where l is the average wavelength, f-number is the ratio of the focal length to the diameter of the transducer, c is the speed of sound and BW is the bandwidth of the transducer. Increasing the lateral resolution can be accomplished by decreasing the average wavelength (increasing the frequency) or the f-number of the transducer. Increasing lateral resolution entails some compromises, decreasing the wavelength of the transducers results in reduced penetration depth, while decreasing the f-number decreases the depth of field. Commercially available nonfocused intravascular ultrasound systems (IVUS) at 40 MHz exhibit better image quality than those of 30 MHz systems [2]. But no commercial systems use focused transducers to enhance lateral resolution. Focusing of small high frequency ultrasound single-element transducers has been accomplished through three methods, the use of a lens, a mirror and shaping the radiating element into a spherical section. The fabrication of miniature lenses has been done through molding [70] and micromilling [71], the fabrication of the mirror was done using conventional machining and press fitting [72]. Shaping the radiating element using a press focusing technique has been successfully demonstrated on a number different piezo

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materials such as PZT, poly(vinylidene fluoride) (PVDF) and lithium niobate [1, 70, 73]. Press focusing does not increase the profile of the device or add additional mechanical components to the transducer, which is an important consideration when trying to simplify the construction of an IVUS catheter. However, brittle materials can be prone to cracking when press focused. 17.3.1

Polymer Transducers

PVDF and its copolymers offer excellent mechanical flexibility for fabricating press focused transducers in addition to having desirable properties for high frequency transducers for medical ultrasound such as wide bandwidth and an acoustical impedance very close to tissue [74]. However, miniaturized polymer transducers have high output impedances and require a high impedance preamplifier in close physical proximity (<1 mm) to prevent reduced signal-to-noise ratio (SNR) due to loading of the transducer from the low impedance cable that attaches the transducer to the corresponding signal processing and imaging console [75, 76]. In IVUS systems, the cable is typically 1.5 m in length, which results in significant signal degradation. Fortunately, PVDF and its copolymers are compatible with silicon integrated circuit (IC) fabrication and lend themselves to microelectromechanical systems (MEMS) techniques to provide transducers in close physical proximity to electronics through the integration of transducers directly on the IC chip. Previous efforts have integrated polymer ultrasonic transducers with electronics [77, 78]. Silicon has been micromachined to improve transducer performance, air backed and epoxy backed transducers fabricated on silicon have been demonstrated [79, 80]. In an integrated version of a transducer chip, the transducer, preamplifier and a 50

buffer driver to transmit the signal from the distal end of the catheter coaxial cable to the IVUS console would be monolithically integrated on a single chip. Protection circuitry to protect the high impedance preamplifier during the transmit pulse is required. Fast overload recovery time, which directly impacts how close to the face of the transducer the system can image, is of critical importance. In the case of IVUS, this would be about 0.75 mm, meaning the amplifier must have recovered from the transmit pulse and be ready to amplify in  1 ms. Ideally, this device would be implemented using a complementary metal oxide semiconductor (CMOS) process (the most common costeffective circuit technology available) instead of specialized high voltage integrated electronics technologies. Finally, the entire chip should fit into a form factor suitable for IVUS implementation. In the following, the fabrication and characterization of micromachined high frequency focused polymer ultrasonic transducers in a manner that is compatible with CMOS microelectronics, and MEMS batch fabrication techniques, are described. The specifics of the electronics are not described here, but the interested reader may find more details elsewhere [75, 76, 81–84]. The transducer is capable of being manufactured on silicon wafers after the completion of CMOS electronics. These two key elements enable the eventual creation of a monolithic transducer chip that does not require modification of the standard circuit fabrication process. This type of transducer chip will likely follow the path of other MEMS devices such as accelerometers, gene chips and digital micromirror arrays, where batch production, high yields and

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economies of scale have lead to relatively low unit costs, potentially making this device inexpensive enough for single use procedures. The techniques presented are scalable to create transducers of arbitrary aperture and focus as well.

17.4 Fabrication Techniques Cross-sectional and plan schematic views of the transducer concepts are shown in Figure 17.3. The plan view illustrates the concept for integration of the polymer transducer with CMOS microelectronics, while the cross-sectional view depicts the critical geometries of the shaped polymer membrane. The shaped polymer membrane constitutes a spherical section whose geometry determines the focal properties of the ultrasonic transducer. The f-number of the device is defined to be r/A, where r is the radius of curvature of the spherical section and A is the diameter of the aperture. The focal point of transducer is at point O. Any change in the center deflection, d, of the membrane for a given aperture changes the radius of curvature, r, and therefore the f-number of the transducer. Circular plates with clamped boundaries deform spherically under uniformly applied differential pressure [85]. However, for a large deflection there is some disagreement as to the exact shape [86–89]; for our purposes, the deformation is spherical, especially away from the edges of the film.

Figure 17.3 Cross-sectional (a) view showing geometric focusing and plan view (b) showing transducer integration with integrated electronics.

The center deflection is a function of membrane geometry, materials properties and differential pressure, as described by Pan [90]: P¼

C1 C2 f ðvÞ E td3 o td þ a4 1  a2

ð17:4Þ

where P is differential pressure, a is radius of the aperture (A/2), E is the Young’s modulus,  is the Poisson ratio, o is the residual stress, d is the center deflection, f() is a dimensionless function equal to 0.957–0.208  for circular membranes, t is the membrane thickness and C1 and C2 are geometrically dependent constants equal to 4.00 and 2.67, respectively for circular membranes. Thus, the f-number for arbitrarily sized apertures can be controlled by changing the differential pressure across the membrane.

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Transducers presented in this paper were fabricated using free-standing 9 mm thick PVDF–TrFE films and 9 mm thick spun-on PVDF–TrFE films. The processes used to fabricate these two types of transducers are essentially the same and, for conciseness, the free-standing technique will be described first in detail. Free-standing 9 mm thick PVDF–TrFE films with a 200 nm thick chromium/gold coating on one side are used as the piezoelectric polymer material (Ktech Corporation, Albuquerque, NM). The process flow is depicted in Figure 17.4. Micromachining is used to form 0.75–2.00 mm diameter apertures in a standard (100) silicon wafer, which is subsequently thermally oxidized to grow a 1.5 mm thick layer of SiO2. The wafer is then diced into 1 cm wide square die to create the substrates for prototype transducers. The die with a 1 cm square polymer film is inverted onto a jig such that the nonmetalized surface of the polymer film is in contact with the silicon die; differential pressure is applied to deform the film into a spherical shape. A conductive epoxy (Tra Con 2907) is then placed on the back of the polymer in the opening to secure the deformed shape of the polymer film, provide backside electrical contact and act as a sound-absorbing backing. A 30 gauge wire is inserted into the epoxy prior to curing to facilitate subsequent electrical circuit connections. The differential pressure on the deformed polymer film is maintained for the duration of epoxy curing. A completed transducer is shown in Figure 17.5 and the schematic of the hybrid electronic circuit is shown in Figure 17.6.

Figure 17.4 Process flow used to fabricate the prototype ultrasonic transducers: (a) a silicon (Si) wafer is micromachined to create an array of holes with diameters between 0.75 and 2.00 mm, after which the wafer is then thermally oxidized to grow a 1.5 m thick SiO2 layer; (b) the wafer is then diced into 1 cm wide square die; (c) the die is laid flat onto a piece of free-standing PVDF film in a jig (the die is now viewed in cross-section through the hole); (d) the die and PVDF film are clamped into the jig against an O-ring forming an air-tight seal, and air pressure is applied to the face of the PVDF film to deflect it into the desired spherical shape; (e) finally, conductive epoxy is injected into the hole and a 30 gage wire is potted into the epoxy; the air pressure is maintained until the epoxy cures, then the transducer chip is removed from the jig.

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Figure 17.5 A prototype MEMS ultrasonic transducer mounted on a custom-printed circuit (PC) board. The 2 mm active element is the circular smooth region in the center. Hybrid electronics (Figure 17.6) are mounted on the back of the PC board.

680 R2 220

4

–

R1

Coax o/p

D3

AD8001

C1

Coax i/p D4

Wire

15pF

3 D1

+

D2

OPAMP

R3 10 k

Transducer

Figure 17.6 Schematic of a hybrid preamplifier circuit mounted on the back of the transducer.

An alternate approach, more compatible with batch fabrication techniques, was also explored. A 9 mm thick PVDF–TrFE polymer film was spun cast onto the oxidized wafer, after which the film was blanket coated with chromiun/gold metallurgy and poled (these processes were obtained from Measurement Specialties Incorporated, Fairfield, NJ). The wafer was then coated top and bottom with photo resist (a photopatternable polymer). The top resist was used as a protective coat and the bottom resist was patterned with circular holes to create a mask to form the apertures on the front of the wafer. The wafer was mounted topside down on a handle wafer then backside etched using deep reactive ion etching (DRIE) to remove the silicon from the back of the wafer until the etching stopped on the thermal oxide on the top. In the next step, the wafer was diced into 11 cm square chips, which were separately processed thereafter. Next, the chip was dipped in hydrofluoric acid to remove the residual oxide left after the DRIE. The chip was then processed identically to the freestanding PVDF films as described previously. Selectivity of photo resist solvents was a significant consideration; it was found that the films were not affected by the use of SVC 175 photo resist stripper (Shipley Marlborough, MA).

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17.5 Testing Methods The test apparatus (Figure 17.7) consists of a pulser (Avtech AVB1–3-C monocycle generator, Avtech Electrosystems, Ogdensburg, NY) to excite the transducer, amplification circuitry and a digitizing oscilloscope (Agilent 54835A, Agilent Technologies, Englewood, CO) to observe signals and collect data. The test setup uses two sets of crossed diode bridges: (i) the limiter to protect the preamplifier from the high voltage pulses ( 50 Vpp) and reduce the preamplifier saturation time; and (ii) the expander to reduce noise contribution from the pulser by isolating it from rest of the apparatus between pulses. Transducer characterization experiments were performed in the pulse echo mode.

Figure 17.7

Schematic of test apparatus used to characterize the fabricated transducers.

The transducers were excited by 40 MHz, 50 Vpp monocycle pulses at a 2 kHz repetition rate in a container of deionized (DI) water to obtain pulse echo responses. The transducer was connected to the pulser by a 0.5 m long, 50 coaxial cable. A glass slide at the bottom of the container provided the reflecting surface. No amplification circuitry was used to minimize parasitic bandwidth components. The axial responses of the transducers were characterized by exciting the transducers as described above, but with an additional 30 dB of amplification in the circuit (Miteq AU-1114, Miteq Corporation, Hauppauge, NY). The transducer was displaced relative to a glass slide and the peak-to-peak voltage of the correspondingpulse echo response was recorded. The lateral resolution of the ultrasonic transducer was determined from radiation pattern measurements using a 30 dB amplifier described above. A precision 60 mm diameter stainless steel wire was used to simulate a point source for back reflected ultrasonic signals from the transducer. The radiation contours of the reflected pulses from the wire were obtained by scanning the transducer laterally at different heights above the wire using a computer controlled XYZ stage with better than 1 mm resolution (Newport Instruments, Irvine, CA) coupled with a 500 MHz 8-bit A/D card (CS 8500 Gage Applied Technologies, Montreal, Canada). Lateral resolution was determined by measuring the radiation beam width at the 6 dB points in the center of the focal region. A silicon MEMS phantom with fixed lines and spaces constructed using DRIE was also imaged. The MEMS phantom

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consisted of sets of fixed lines and spaces micromachined in a silicon wafer ranging from 160 mm lines and spaces to 20 mm lines and spaces. The top of the silcon wafer is polished and the side walls are nearly vertical. This provides flat highly smooth reflecting surfaces and an abrupt boundary between the reflecting surface and the space. Tissue images were collected using a high impedance preamplifier, as tissue is less echogenic than glass, and, consequently, a preamplifier was needed to maintain adequate SNR. The transducer was mounted onto a PC board incorporating the 12 dB gain preamplifier, as shown in Figures 17.5 and 17.6. A sample of human cadaveric aorta was filleted and pinned flat onto a paraffin base. The aorta was covered with DI water and the vertical height of the transducer relative to the aorta was adjusted to obtain the maximum pulse echo response from the tissue, thereby placing the focal point of the transducer in the tissue. The transducer was then scanned using a motorized stage in one direction in increments of 0.05 mm. The reflection at each position was sampled by the oscilloscope at 1 GSa/s and saved as a file to create one scan line. The scan lines were subsequently compiled and standard imaging software (MATLAB, The Mathworks Inc., Natick, MA,) was used to create the 256 level gray scale image of the aorta.

17.6 Results A typical pulse echo response, obtained from a transducer with a 2.00 mm aperture is shown in Figure 17.8. The pulse echo exhibits minimal ringing, which indicates a relatively large bandwidth. The corresponding power spectrum is shown in Figure 17.9. The axial resolution of the transducer was determined from the full-width-at-half-maximum (FWHM) of the pulse echo envelope, which was obtained via a Hilbert Transform of the pulse echo response. The FWHM has a value of 28 ns, which results in an axial resolution 50 40

Voltage (milli Volts)

30 20 10 0 –10 –20 –30 5.9

6

6.1 6.2 6.3 Time (micro s)

6.4

6.5

Figure 17.8 Pulse echo response of a 2 mm transducer shows minimal ringing. Time is referenced from the application of the excitation pulse.

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Power spectral density (dB)

10 0 –10 –20 –30 –40 –50 –60

0

20

40 60 Frequency (MHz)

80

100

Figure 17.9 Pulse power spectrum of the transducer from Figure 17.8, showing a center frequency of 36 MHz and a 6 dB bandwidth of 120 %.

of 21 mm, assuming a sound velocity of 1500 m/s in water. The transducer characterized here has a center frequency of 36 MHz and a 6 dB bandwidth of 120 %. Other transducers that were fabricated using the same technique and material exhibited center frequencies of 30–45 MHz and 6 dB bandwidths of 80–110 %. The variation in characteristics of the various transducers is attributed to nonuniformity in polymer film thickness and epoxy curing. These results compare favorably with those from spherically focused PVDF transducers that are fabricated using press focusing techniques [1, 74, 91]. The results for three transducers with 2.00 mm diameter apertures and membrane center deflections of 172, 123 and 102 mm – 4 mm, respectively, are shown in Figure 17.10. These center deflections should result in f-numbers of 1.50, 2.00 and 2.45, respectively; they are

f = 1.47 f = 2.08 f = 2.41

Normalized Amplitude

1.2 H

1

H H

0.8 0.6 0.4 0.2 0

2

4 6 Distance (mm)

8

Figure 17.10 Axial radiation patterns for three, 2 mm aperture transducers with center deflection of 172, 123 and 102 – 4 m, showing that observed focal lengths are in good agreement with predicted values based upon measured center deflections.

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Normalized Amplitude

in general agreement with the values of 1.47, 2.08 and 2.41, which are derived from the axial response measurements (the error bars in the figure show the variation expected due to the error in measurement of center deflection). Further range in f-number is possible by varying the pressure even more; Figure 17.11 shows the effect of varying pressure between 20 and 5 psi resulting in f-numbers between 2 and 4.15 with a corresponding increase in focal point and depth of field as expected. 1.2 1 0.8 0.6 0.4 0.2 0 2 3 4 5 6 7 8 9 10 11 12 13 14 Distance (mm)

Figure 17.11 Axial radiation patter for 2 mm aperture transducers where the four numbers are changed from 2 to 4.15 by varying the pressure used to deflect the polymer film.

The continuous wave axial radiation pattern of a 2 mm diameter transducer with geometric focus at 7.1 mm was obtained by exciting the transducer with a gated sinusoidal pulse having 15 cycles and using a 20 mm stainless steel wire as a target. The continuous wave axial radiation pattern for an identical spherical transducer was simulated in MATLAB using the expressions given by H.T. O’Neil [92]. Figure 17.12 shows good

Figure 17.12 Comparison of experimental with simulated axial radiation patterns of a 2 mm transducer showing good agreement with a Pearson correlation of 0.99.

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agreement between the axial experimental and simulated radiation patterns with a Pearson correlation coefficient of 0.99, over the axial range from 5 to 10 mm, indicating that the fabricated transducers are spherical. Deviations from the model are attributed primarily to the use of a wire reflector instead of a point source. Radiation patterns are shown in Figure 17.13 for transducers with different apertures and f-numbers. Radiation patterns are in good agreement with models obtained using the methods described by Arditi for spherically focused transducers [93].

Figure 17.13 Radiation patterns for (a) a 2 mm aperture F-2.5 transducer; (b) a 1 mm aperture F-3.1 transducer; (c) a 0.75 mm F-3.3 transducer; and (d) experimental and predicted results for from the transducer shown in (c) for the subset or radiation contours for –3, –6 and –9 dB.

Imaging resolution was demonstrated by imaging a silicon micromachined phantom consisting of fixed beams and spaces ranging from 20 to 160 mm as shown in Figure 17.14. It can be seen that the transducer can differentiate line-space pairs as narrow as 60 mm and some structures are observable even at 40 mm.

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(a)

(b) 20 40 60

80

100

120

140

160 (c)

Figure 17.14 (a) Electron microscope pictures of micromachined silicon phantom; (b) schematic of lines and spaces on the phantom; and (c) gray scale image of phantom obtained with a 1 mm transducer showing good lateral resolution up to 60 mm and some structures even at 40 mm.

Tissue was imaged as described earlier using a 2.00 mm diameter aperture and an f-number of 2.08 transducer mounted onto a PC board incorporating the preamplifier. The close physical proximity ( 2 mm) of the transducer to the preamplifier negated the effect of  4 dB loss in signal through the 0.5 m-long coaxial cable used to connect the transducer to the Miteq 30 dB amplifier. The overall system was found to have an active gain of 42 dB. An image of the aorta is depicted in Figure 17.15. This image clearly shows the subsurface elastin bands of the aortic tissue as well as the top and bottom surfaces. Lastly, Figure 17.16 shows a comparison of a 1 mm focused PVDf–TrFE transducer with two commercially available systems and about the same level of cadaveric human coronary artery. The performance of the micromachined ultrasonic transducer compares very favorably with previously published work based on conventional methods [2].

(a)

(b)

1 mm

Figure 17.15 (a) Ultrasound image of a human cadaveric aorta obtained using a 2 mm focused polymer transducer; (b) corresponding histological image.

Miniature High Frequency Focused Ultrasonic Transducers (c)

(b)

(a)

1 mm

351

1 mm

1 mm

Figure 17.16 Images of cadaveric coronary artery obtained from the same cadaveric artery from about the same with: (a) a Volcano Revolution IVUS catheter; (b) MEMS focused transducer; and (c) a Volcano Eagle Eye IVUS catheter.

17.7 Conclusion The design, fabrication and characterization of novel micromachined focused ultrasonic transducers for minimally invasive medical imaging procedures have been demonstrated. Transducers were fabricated using a membrane deflection technique to produce spherical sections in a piezoelectric polymer film. Pulse echo responses showed minimal ringing and wide bandwidths characteristics of 80–110 %. Axial resolution of 50 mm and lateral resolution of 51–92 mm were achieved. A preamplifier circuit incorporated into a hybrid package with the ultrasonic transducer exhibited thermal noise of 5.3 nV/Hz½. The transducers were used to image human cadaveric aorta to reveal high-resolution subsurface structures. The techniques and results presented in this paper prove the feasibility of a CMOS compatible fabrication of focused polymer transducers for minimally invasive ultrasonic imaging. The success of this approach justifies investment into additional development towards the integration of CMOS microelectronics with focused polymeric transducers.

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7. Cook, S., Wenaweser, P., Togni, M., et al. (2007) Incomplete Stent Apposition and Very Late Stent Thrombosis After Drug-Eluting Stent Implantation, Circulation, 115, 2426–34. 8. Mehta, S. K., McCrary, J. R., Frutkin, A. D., et al. (2007) Intravascular ultrasound radiofrequency analysis of coronary atherosclerosis: an emerging technology for the assessment of vulnerable plaque, Eur. Heart J., 28, 1283–8. 9. Siqueira, D. A., Abizaid, A. A., Costa, J. d. R., et al. (2007) Late incomplete apposition after drug-eluting stent implantation: incidence and potential for adverse clinical outcomes, Eur. Heart J., 28, 1304–9. 10. Feres, F., Costa, J. R. and Abizaid, A. (2006) Very late thrombosis after drug-eluting stents, Catheterization and Cardiovascular Interventions, 68, 83–8. 11. Mintz, G. S. and Weissman, N. J. (2006) Intravascular Ultrasound in the Drug-Eluting Stent Era, J. Am. Coll. Cardiol., 48, 421. 12. De Scheerder, I., De Man, F., Herregods, M. C., et al. (1993) The use of intracoronary ultrasound for quantitative assessment of coronary artery lumen diameter and for on-line evaluation of angioplasty results, Acta Cardiol., 48, 171–81. 13. Cavaye, D., White, R., Kopchok, G., et al. (1992) Three-dimensional intravascular ultrasound imaging of normal and diseased canine and human arteries, J. Vasc. Surg., 16, 509–17. 14. Siegel, R., Ariani, M., Fishbein, M., et al. (1991) Histopathologic validation of angioscopy and intravascular ultrasound, Circulation, 84, 109–17. 15. Yoshida, K., Yoshikawa, J., Akasaka, T., et al. (1992) Intravascular ultrasound imaging – in vitro and vivo validation., Jpn Circ. J., 56, 572–7. 16. Ng, K. H., Evans, J. L., Vonesh, M. J., et al. (1994) Arterial imaging with a new forward-viewing intravascular ultrasound catheter, II. Three-dimensional reconstruction and display of data, Circulation, 89, 718–23. 17. Gussenhoven, E. J., Frietman, P. A., The, S. H., et al. (1991) Assessment of medial thinning in atherosclerosis by intravascular ultrasound., Am. J. Cardiol., 68, 1625–32. 18. von Birgelen, C., Di Mario, C., Li, W., et al. (1996) Morphometric analysis in three-dimensional intracoronary ultrasound: an in vitro and in vivo study performed with a novel system for the contour detection of lumen and plaque, Am. Heart J., 132, 516–27. 19. Shekhar, R., Cothren, R. M., Vince, D. G., et al. (1999) Three dimensional segmentation of luminal and adventitial borders in serial intravascular ultrasound images, Computerized Medical Imaging and Graphics, 23, 299–309. 20. Wenguang, L., Bosch, J. G., Zhong, Y., et al. (1991) Semiautomatic frame-to-frame tracking of the luminal border from intravascular ultrasound, Proceedings of Computers in Cardiology Venice, Italy, 353–6. 21. Zhang, X., McKay, C., and Sonka, M. (1998) Tissue characterization in intravascular ultrasound images., IEEE Trans. Med. Imaging, 17, 889–99. 22. von Birgelen, C., Di Mario, C., Reimers, B., et al. (1996) Three-dimensional intracoronary ultrasound imaging. Methodology and clinical relevance for the assessment of coronary arteries and bypass grafts., J Cardiovasc. Surg. (Torino), 37, 129–39. 23. Frank, R. J., McPherson, D. D., Chandran, K. B. and Dove E. L. (1996) Optimal surface detection in intravascular ultrasound using multi-dimensional graph search, Proceedings of Computers in Cardiology 45–8. 24. Sonka, M., Liang, W., Zhang, X., et al. (1995) Three-dimensional automated segmentation of coronary wall and plaque from intravascular ultrasound pullback sequences, Proceedings of Computers in Cardiology Vienna, Austria, 637–40. 25. Nair, A., Kuban, B. D., Obuchowski, N. and Vince, D. G. (2001) Assessing spectral algorithms to predict atherosclerotic plaque composition with normalized and raw intravascular ultrasound data, Ultrasound in Medicine and Biology, 27, 1319–31.

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26. Nair, A., Kuban, B., Tuczu, E. M., et al. (2002) Coronary Plaque Classification With Intravascular Ultrasound Radiofrequency Data Analysis, Circulation, 106, 2200–6. 27. de Korte, C. L. and van der Steen, A. F. W. (2002) Intravascular ultrasound elastography: an overview, Ultrasonics, 40, 859. 28. Kawase, Y., Hoshino, K., Yoneyama, R., et al. (2005) In vivo volumetric analysis of coronary stent using optical coherence tomography with a novel balloon occlusion-flushing catheter: A comparison with intravascular ultrasound, Ultrasound in Medicine & Biology, 31, 1343. 29. Personal Communication (2007) Vince G., Vice President of Clinical Affairs and Advanced R&D, Volcano Corporation, Rancho Cordova, CA. 30. Costa, M. A., Angiolillo, D. J., Tannenbaum, M., et al. (2008) Impact of Stent Deployment Procedural Factors on Long-Term Effectiveness and Safety of Sirolimus-Eluting Stents (Final Results of the Multicenter Prospective STLLR Trial), Am. J. Cardiol., 101, 1704–11. 31. Rosamond, W., Flegal, K., Furie, K., et al. (2008) Heart Disease and Stroke Statistics – 2008 Update: A Report From the American Heart Association Statistics Committee and Stroke Statistics Subcommittee, Circulation, 117, 125–46. 32. Falk, E., Shah, P., and Fuster, V. (1995) Coronary plaque disruption, Circulation, 92, 657–71. 33. Madjid, M., Zarrabi, A., Litovsky, S., et al. (2004) Finding Vulnerable Atherosclerotic Plaques: Is It Worth the Effort?, Arterioscler. Thromb. Vasc. Biol., 24, 1775–82. 34. Virmani, R., Burke, A., Farb, A. and Koldgie, F. (2002) Pathology of the Unstable Plaque, Prog. Cardiovasc. Diseases, 44, 349–56. 35. Qiao, J. H. and Fishbein, M. C. (1991) The severity of coronary atherosclerosis at sites of plaque rupture with occlusive thrombosis, J. Am. Coll. Cardiol., 17, 1138–42. 36. Davies, M. J. and Thomas, A. (1984) Thrombosis and acute coronary-artery lesions in sudden cardiac ischemic death, New Eng. J. Med., 310, 1137–40. 37. Falk, E. (1983) Plaque rupture with severe pre-existing stenosis precipitating coronary thrombosis. Characteristics of coronary atherosclerotic plaques underlying fatal occlusive thrombi, Brit. Heart J., 50, 127–34. 38. Horie, T., Sekiguchi, M. and Hirosawa, K. (1978) Coronary thrombosis in pathogenesis of acute myocardial infarction. Histopathological study of coronary arteries in 108 necropsied cases using serial section, Brit. Heart J., 40, 153–61. 39. Giroud, D., Li, J. M., Urban, P., et al. (1992) Relation of the site of acute myocardial infarction to the most severe coronary arterial stenosis at prior angiography [see comments], Am. J. Cardiol., 69, 729–32. 40. Little, W. C., Constantinescu, M., Applegate, R. J., et al. (1988) Can coronary angiography predict the site of a subsequent myocardial infarction in patients with mild-to-moderate coronary artery disease?, Circulation, 78, 1157–66. 41. Ambrose, J. A., Tannenbaum, M. A., Alexopoulos, D., et al. (1988) Angiographic progression of coronary artery disease and the development of myocardial infarction, J. Am. Coll. Cardiol., 12, 56–62. 42. Hermiller, J. B., Tenaglia, A. N., Kisslo, K. B., et al. (1993) In vivo validation of compensatory enlargement of atherosclerotic coronary arteries, Am. J. Cardiol., 71, 665–8. 43. Glagov, S., Weisenberg, E., Zarins, C. K., et al. (1987) Compensatory enlargement of human atherosclerotic coronary arteries, New Eng. J. Med., 316, 1371–5. 44. Klingensmith, J., Vince, D., Kuban, B., et al. (2000) Assessment of Coronary Compensatory Enlargement by Three-dimensional Intravascular Ultrasound, International Journal of Cardiac Imaging, Int. J. Cardiac Imaging, 16, 87–98. 45. Schoenhagen, P., Ziada, K. M., Kapadia, S. R., et al. (2000) Extent and direction of arterial remodeling in stable versus unstable coronary syndromes: an intravascular ultrasound study, Circulation, 101, 598–603.

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46. Ge, J., Erbel, R., Gerber, T., et al. (1994) Intravascular ultrasound imaging of angiographically normal coronary arteries: a prospective study in vivo, Brit. Heart J., 71, 572–8. 47. Hodgson, J., Reddy, K., Suneja, R., et al. (1993) Intracoronary ultrasound imaging: correlation of plaque morphology with angiography, clinical syndrome and procedural results in patients undergoing coronary angioplasty, J. Am. Coll. Cardiol., 21, 35–44. 48. Nissen, S. E., Gurley, J. C., Grines, C. L., et al. (1991) Intravascular ultrasound assessment of lumen size and wall morphology in normal subjects and patients with coronary artery disease, Circulation, 84, 1087–99. 49. Tobis, J. M., Mallery, J., Mahon, D., et al. (1991) Intravascular ultrasound imaging of human coronary arteries in vivo. Analysis of tissue characterizations with comparison to in vitro histological specimens, Circulation, 83, 913–26. 50. Evans, J. L., Ng, K. H., Wiet, S. G., et al. (1996) Accurate three-dimensional reconstruction of intravascular ultrasound data. Spatially correct three-dimensional reconstructions, Circulation, 93, 567–76. 51. Sangiorgi, G. M., Clementi, F., Cola, C. and Biondi-Zoccai, G. (2007) Plaque vulnerability and related coronary event prediction by intravascular ultrasound with virtual histology: ldquoIt’s a long way to tipperaryrdquo?, Catheterization and Cardiovascular Interventions, 70, 203–10. 52. Iakovou, I., Schmidt, T., Bonizzoni, E., et al. (2005) Incidence, Predictors, and Outcome of Thrombosis After Successful Implantation of Drug-Eluting Stents, JAMA, 293, 2126–30. 53. Kuchulakanti, P. K., Chu, W. W., Torguson, R., et al. (2006) Correlates and Long-Term Outcomes of Angiographically Proven Stent Thrombosis With Sirolimus- and PaclitaxelEluting Stents, Circulation, 113, 1108–13. 54. Ong, A. T. L., Hoye, A., Aoki, J., et al. (2005) Thirty-day incidence and six-month clinical outcome of thrombotic stent occlusion after bare-metal, sirolimus, or paclitaxel stent implantation, J. Am. Coll. Cardiol., 45, 947–53. 55. Jeremias, A., Sylvia, B., Bridges, J., et al. (2004) Stent Thrombosis After Successful SirolimusEluting Stent Implantation, Circulation, 109, 1930–2. 56. Park, D.-W., Hong, M.-K., Mintz, G. S., et al. (2006) Two-Year Follow-Up of the Quantitative Angiographic and Volumetric Intravascular Ultrasound Analysis After Nonpolymeric Paclitaxel-Eluting Stent Implantation: Late ‘‘Catch-Up’’ Phenomenon From ASPECT Study, J. Am. Coll. Cardiol., 48, 2432–9. 57. Moreno, R., Fernandez, C., and Macaya, C. (2006) Drug-eluting stent thrombosis: Results from a pooled analysis including 10 randomized studies: Reply, J. Am. Coll. Cardiol., 47, 214–15. 58. Bavry, A. A., Kumbhani, D. J., Helton, T. J. and Bhatt, D. L. (2005) What is the risk of stent thrombosis associated with the use of paclitaxel-eluting stents for percutaneous coronary intervention?: A meta-analysis, J. Am. Coll. Cardiol., 45, 941–6. 59. Bavry, Kumbhani, Helton, and Bhatt, (2005) Risk of Thrombosis With the Use of SirolimusEluting Stents for Percutaneous Coronary Intervention (from Registry and Clinical Trial Data), Am. J. Cardiol., 95, p. 1469. 60. Regar, E., Lemos, P. A., Saia, F., et al. (2004) Incidence of thrombotic stent occlusion during the first three months after sirolimus-eluting stent implantation in 500 consecutive patients, Am. J. Cardiol., 93, p. 1271. 61. Williams, D. O., Abbott, J. D. and Kip, K. E. (2006) Outcomes of 6906 Patients Undergoing Percutaneous Coronary Intervention in the Era of Drug-Eluting Stents: Report of the DEScover Registry, Circulation, 114, 2154–62. 62. Urban, P., Gershlick, A. H., Guagliumi, G., et al. (2006) Safety of Coronary Sirolimus-Eluting Stents in Daily Clinical Practice: One-Year Follow-Up of the e-Cypher Registry, Circulation, 113, 1434–41.

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63. Ako, J., Morino, Y., Honda, Y., et al. (2005) Late Incomplete Stent Apposition After SirolimusEluting Stent Implantation: A Serial Intravascular Ultrasound Analysis, J. Am. Coll. Cardiol., 46, 1002–5. 64. Mintz, G. S., Shah, V. M. and Weissman, N. J. (2003) Regional Remodeling as the Cause of Late Stent Malapposition, Circulation, 107, 2660–3. 65. Hong, M.-K., Mintz, G. S., Lee, C. W., et al. (2004) Incidence, Mechanism, Predictors, and Long-Term Prognosis of Late Stent Malapposition After Bare-Metal Stent Implantation, Circulation, 109, 881–6. 66. Matsumoto, D., Shite, J., Shinke, T., et al. (2007) Neointimal coverage of sirolimus-eluting stents at 6-month follow-up: evaluated by optical coherence tomography, Eur. Heart J., 28, 961–7. 67. de Smet, B. J. G. L. and Zijlstra, F. (2007) A look at drug eluting stents with optical coherence tomography, Eur. Heart J., 28, 918–9. 68. Kume, T., Akasaka, T., Kawamoto, T., et al. (2006) Assessment of Coronary Arterial Plaque by Optical Coherence Tomography, Am. J. Cardiol., 97, p. 1172. 69. Burke, A. and Virmani, R. (2007) Pathophysiology of acute myocardial infarction., Med. Clin. North Am., 91, 553–72. 70. Cannata, J. M., Ritter, T. A., Chen, W.-H., et al. (2003) Design of efficient, broadband singleelement (20–80 MHz) ultrasonic Transducers for medical imaging applications, IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, 50, 1548–57. 71. Li, J., Friedrich, C. R. and Keynton, R. S. (2002) Design And Fabrication Of A Miniaturized, Integrated, High-Frequency Acoustical Lens-Transducer System, J. Micromech. Microeng., 12, 219–28. 72. Lockwood, G. R., Ryan, L. K. and Foster, F. S. (1993) A 45 to 35 MHz Needle-Based Ultrasound System for Invasive Imaging, Ultrasonic Imaging, 15, 1–13. 73. Lockwood, G. R., Turnbull, D. H. and Foster, F. S. (1993) Fabrication of high frequency spherically shaped ceramic transducers, IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, 41, 231–5. 74. Foster, F. S., Harasiewicz, K. A. and Sherar, M. S. (2000) A History of Medical and Biological Imaging with Polyvinylidene Fluoride (PVDF) Transducers, IEEE Transactions on Ultrasonics, Ferroelectrics and Frequency Control, 47, 1363–71. 75. Lockwood G. R. and Hazard C. R. (1997) Development of small aperture polymer transducers for high frequency imaging, Proceeding of the 1997 Ultrasonics Symposium, Toronto, Canada, 5 October 1997. 76. Lockwood, G. R. and Hazard, C. R. (1998) Miniature polymer transducers for high frequency medical imaging, SPIE International Symposium on Medical Imaging 1998, 3341, 228–36. 77. Swartz, R. G. and Plummer, J. D. (1979) Integrated Silico-PVF2 Acoustic Transducer Arrays, IEEE Transactions on Electron Devices, 26, 1921–31. 78. Fiorillo, A. S., Spiegel, J. V. D., Bloomfield, P. E. and Esmail-Zandi, D. (1990) A P(VDF-TrFE)based Integrated Ultrasonic Transducer, Sensors and Actuators A, 22, 719–25. 79. Mo, J., Robinson, A. L., Fitting, D. W., et al. (1990) Micromachining for Improvement of Integrated Ultrasonic Transducer Sensitivity, IEEE Transactions on Electron Devices, 37, 134–9. 80. Sleva, M. Z., Briggs, R. D. and Hunt, W. D. (1996) A Micromachined Poly(vinylidene fluoridetrifluroethylene) Transducer for Pulse-Echo Ultrasound Applications, IEEE Transactions on Ultrasonics, Ferroelectrics and Frequency Control, 43, 257–62. 81. Fuller, M. I., Blalock, T. N., Hossack, J. A. and Walker, W. F. (2007) Novel transmit protection scheme for ultrasound systems, IEEE Transactions on Ultrasonics, Ferroelectrics and Frequency Control, 54, 79. 82. Fan J., Talman J., Fleischman A. and Garverick S. L. (2004) Integrated Amplifier with Active Limiter for Intravascular Ultrasonic Imaging, IASTED International Conference on Circuits, Signals, and Systems, Clearwater, FL, USA, 433–8.

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83. Lockwood, G. R., Hunt, J. W. and Foster, F. S. (1991) The Design of Protection Circuitry for High-Frequency Ultrasound Imaging Systems, IEEE Transactions On Ultrasonics, Ferroelectrics and Frequency Control, 38, 48–55. 84. Talman, J. R. and Garverick, S. L. (2001) Integrated-Circuit Implementation of a Matched-Cell Dynamic Focusing Architecture for a 5-Channel, 50-Mhz Planar Annular Array, Proceedings of the 2001 IEEE Ultrasonics Symposium, 2, 1109–1112. 85. Timoshenko, S. and Woinowsky-Krieger, S., Theory of plates and shells, McGraw-Hill, New York, 1987. 86. Shlepak, M. and Dugundji, J. (1998) Large Deflections of Clamped Circular Plates Under Initial Tension and Transitions to Membrane Behavior, J. App. Mech.s, 65, 107–15. 87. Small, M. K. and Nix, W. D. (1992) Analysis of the accuracy of the bulge test in determining the mechanical properties of thin films, J. Mat. Res., 7, 1553–63. 88. Marker, D. K. and Jenkins, C. H. (1997) Surface precision of optical membranes with curvature, Optics Express, 1, 324–31. 89. Eaton, W. P., Bitsie, F., Smith, J. H. and Plummer, D. W. (1999) A New Analytical Solution for Diaphragm Deflection and its Application to a Surface-Micromachined Pressure Sensor, in Conference on Modeling and Simulation of Microsystems, San Juan Marriott Resort & Stellaris Casino, San Juan, Puerto Rico, 6 April 1999, 640–3. 90. Pan, J. Y., A Study of Suspended-Membrane and Acoustic Techniques for the Determination of the Mechanical Properties of Thin Polymer Films, PhD Thesis, Massachusetts Institute of Technology, Cambridge, MA, 1991. 91. Sherar, M. D. and Foster, F. S. T. (1989) The design and fabrication of high frequency poly(vinylidene fluoride) transducers, Ultrasonic Imaging, 11, 75–94. 92. H. T. O’Neil, (1949) Theory of Focusing Radiators, JASA, 21, 516–26. 93. Arditi, M., Foster, F. S. and Hunt, J. W. (1981) Transient fields of concave annular arrays, Ultrasonic Imaging, 13, 37–61.

18 Catheters for Thrombosis Sample Exfoliation in Blood Vessels Using Piezoelectric Polymer Fibers Yoshiro Tajitsu Smart Structures and Materials Laboratory, Department of Electrical Engineering, Graduate School of Engineering, Kansai University, Japan

18.1 Introduction A crystalline polymer is an extremely large molecular aggregate of chains whose individual units are covalently bound to each other [1, 2]. That is, it has a complex high-order structure with intermingled crystalline and amorphous regions. Crystalline polymer films of 100 % crystallinity cannot be obtained by traditional methods and amorphous components are always included to obtain certain complex high-order structures [1, 2]. Furthermore, no direct relationship has been found between the macro-optical properties and crystal characteristics of these films [1, 2]. When first reported about thirty years ago in Japan, the ferroelectricity of the copolymer of poly(vinylidene fluoride) and trifluoroethylene (PVDF–TrFE)), which is a crystalline polymer, came as a surprise [3, 4]. At present, PVDF–TrFE is a well known example of a ferroelectric polymer [1–5]. Also, this copolymer and poly(vinylidene fluoride) (PVDF) exhibit a tensile piezoelectric constant [1–5], which has been found to continuously increase with the increasing degree of crystallinity and orientation of crystallites in a polymer membrane over these past thirty years [1–5]. However, the practical application of these piezoelectric polymers is limited, because the piezoelectric constant of a ferroelectric polymer membrane is lower than that of inorganic piezoelectric materials [1–5]. Also, ferroelectric polymers such as PVDF

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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require poling for the realization of piezoelectricity in the entire polymer membrane [1–5]. However, the poling of a ferroelectric polymer membrane is very difficult because it requires a high electric field, almost the same as the breakdown strength of the ferroelectric polymer membrane [1–5]. On the other hand, there is one type of polymer that shows shear piezoelectricity, namely, a chiral polymer, in which molecules with asymmetric carbon atoms have a helical orientation [1, 2, 6]. A chiral polymer requires no poling treatment but a drawing treatment to realize piezoelectricity in its entire polymer membrane [1, 2, 4, 6]. However, -helix polypeptides, which are a type of chiral polymer, show a shear piezoelectric constant lower than 1 pC/N [1, 2, 4, 6], which is very small. Shear piezoelectricity can be understood at the molecular level as follows [1, 2, 4]. In general, shear piezoelectricity originates from a dipole accompanying an asymmetric carbon. Shear stress is applied to molecular chains of -helix polypeptides through the side chain. This time, all the atoms that constitute the molecular chain are displaced. The effective motion in -helix polypeptide that induces a change in the polarization is caused by the rotation of the plane on which the C–N bond and C¼O are, round the C–N bond. As a result, the change in the polarization is due to applied stress, that is, piezoelectricity in polypeptides. It seemed that the piezoelectric constant is small because the degree of orientation of chiral polymers is low [1, 2, 4, 6]. Therefore, it was speculated that a chiral polymer membrane would show a large piezoelectric constant if a highly orientated chiral polymer could be developed. On the other hand, recently, poly-l-lactic acid (PLLA), a chiral polymer, has attracted attention as a polymer that exhibits a high shear piezoelectric constant [1, 2, 4, 6]. In this chapter, we focus on a PLLA fiber that is a model chiral polymer fiber, and report on the measurement of its basic piezoelectric characteristics under the application of DC and AC voltages. Finally, from the results, as one example of a proposed application model in biomedical engineering [7], it was shown that a PLLA fiber can remove objects such as an electric control catheter and tweezers.

18.2 Piezoelectricity of Polymer Film and Fiber In polymer film, amorphous components are always present in complex high-order structures. For polymer films and fibers, macroscopic symmetry must be considered on the basis of the point group theory. As seen in Figure 18.1 [8, 9], asymmetry is imparted to the film by conventional methods such as drawing and poling [1, 2, 4, 8]. The purpose of drawing is to arrange chain molecules in a polymer film along one direction of the entire film, macroscopically [1, 2, 10]. The purpose of poling is to arrange the dipole moments of molecules along one direction of the entire film, macroscopically. Actually, the piezoelectricity of an isotropic film does not occur even though the macroscopic piezoelectric d-constant (dijm, i ¼ 1 3, j ¼ 1 6) exists in the crystal state [1, 2, 10]. The point group of drawn and poled films of ferroelectric polymers such as PVDF is C2v, as shown in Figure 18.1 [8, 9]. The symmetry is the same as that of the PVDF crystal. Drawn and poled PVDF films exhibit high tensile piezoelectricity, macroscopically [1, 2, 10].

Catheters for Thrombosis Sample Exfoliation in Blood Vessels

359

D2 0 0 0 1

disappearance of vertical 2 mirror plane

1 drawing

3 drawing

0 0 0

0 0 d14 0 0 0 0 0 d25 0 0 0 0 0 d36 PLLA crystal m 0 0 d14 0 0 0 0 0 –d14m 0 0 0 0 0 0

PLLA D∞

chiral molecules

3 2 vertical mirror plane 0 dij = 0 0

0 0 0 0 0 0 0 0 0

3 poling

0 0 0 0 0 0

1 2

D∞v

drawing

m 0 0 0 d14 0 0 0 0 0 0 d25m 0 m m d31m d32 d33m 0 0 d36

PVDF C2v 0 0 0 d14 0 0 0 0 0 0 d25 0 d31 d32 d33 0 0 d36 PVDF crystal

Figure 18.1 Macro-piezoelectricity imparted by drawing and poling treatment of chiral polymer films.

The point group of a drawn polymer film is D1v [1, 2, 4, 6, 10], as shown in Figure 18.1 [8, 9]. In this case, it is important that a mirror plane perpendicular to the film surface exists. Furthermore, in this case, no piezoelectricity arises. However, the PLLA molecule has chirality, as shown in Figure 18.2 [8, 9]. The crystal structure is based on a base-centered orthorhombic unit cell [11, 12]. Here, a, b and c are the lengths of the unit cell. It contains two 10/3 helical chains arranged along the c-axis, and the point group is D2. PLLA is a chiral polymer and, in the crystal, the chain molecules form a helical structure. The crystal structure of PLLA is characterized by this helical structure. That is, the piezoelectric characteristics of the PLLA crystal are governed by this helical structure. In the PLLA crystal, three piezoelectric d-constants, d14, d25 and d36, as shown in Figure 18.2 [8, 9], are present. Piezoelectricity is generated macroscopically by the drawing of the PLLA film, as shown in Figure 18.1. In this case, the mirror plane disappears and the point group becomes D1. The macropiezoelectric constants d14m and d25m ¼ d14m of an oriented PLLA film are based on d14, d25 and d36 of the crystal state [1, 2, 6, 10]. However, the notation ‘m’ of d14m and d25m is very complicated. Hereafter, d14 as d14m and d25 as d25m are used.

360

Biomedical Applications of Electroactive Polymer Actuators

Figure 18.2 Crystal structure of poly-l-lactic acid (PLLA).

In considering the PLLA fiber, the shear strain sm (m ¼ 4, 5) induced by the electric field Ej (j ¼ 1, 2) is represented by the equations below. The charge Dj (j ¼ 1, 2) induced by shear stress Tm (m ¼ 4, 5) is also represented [1, 2, 10]: D1 ¼ d14 T4

ð18:1Þ

D2 ¼ d25 T5

ð18:2Þ

s4 ¼ d14 E1

ð18:3Þ

s5 ¼ d25 E2

ð18:4Þ

Referring to Equations (18.3) and (18.4), the induced shear strain is perpendicular to the direction of the electric field. In the same way, the direction of induced charge is perpendicular to the direction of stress.

Catheters for Thrombosis Sample Exfoliation in Blood Vessels

361

18.3 Simple Measurement Method for the Bending Motion of Piezoelectric Polymer Fiber PLLA fibers with various drawing ratios were prepared according to a traditional fiber preparation process [13], and their piezoelectricity was evaluated. For example, a PLLA fiber with a molecular weight of 10 000, which was drawn three times at 70 °C, had a piezoelectric modulus of 0.5 pC/N [14]. A PLLA fiber with a molecular weight of 200,000, which was drawn seven times at 90 °C, showed the largest piezoelectric modulus of 25 pC/N [14]. For determining the bending motion due to the piezoelectricity of a PLLA fiber, one end of the PLLA fiber was clamped, as shown in Figure 18.3 [8, 15, 16], alternating strain was applied to the PLLA fiber using a vibrator and the response stress of the fiber was detected using a load cell. Simultaneously, the response charge due to the shear piezoelectric effect of the PLLA fiber was detected using a charge amplifier. That is,

charge amplifier stress

charge

A/D

clamped

free end

personal computer

strain polymer fiber

Figure 18.3

Schematic of simple system for measuring piezoelectricity of fiber. Charge signal (V)

2 1 0 –1

Stress signal (V)

–2 10 0 –10 0.0

0.2

0.4 0.6 Time (s)

0.8

1.0

Figure 18.4 Stress and charge signals generated by the bending motion of PLLA fiber.

362

Biomedical Applications of Electroactive Polymer Actuators

response stress and charge with a frequency equal to that of the applied alternating strain were detected, and the piezoelectric modulus was evaluated from the ratio of the response charge signal to the response stress signal. The piezoelectric modulus of the PLLA fiber was thus evaluated using this measuring system. A typical experimental result is shown in Figure 18.4 [14], in which the response stress and response charge are shown. In this case, the frequency of the alternating strain is 1Hz. The period of the applied alternating strain is equal to those of the response stress and charge.

18.4 Piezoelectric Motion of Poly-L-Lactic Acid (PLLA) Fiber To drive the PLLA fiber on the basis of Equations (18.3) and (18.4), an effective method of applying an electric field perpendicular to the PLLA fiber axis must be devised. Some cases were introduced here in which the motion of the free end was observed using a charge coupled device (CCD) camera when a voltage was applied to the PLLA fiber, as shown in Figure 18.5 [8, 9, 14–16]. One end of the PLLA fiber was fixed and the other end of the fiber was left free. The length from the free end to the fixed position was about 12 cm. As a result, the electric field causes shear strain due to the piezoelectricity of the PLLA fiber.

Figure 18.5 Schematic of the system for measuring the bending motion of PLLA fiber when voltage is applied.

First, in the application of an AC applied voltage of 500 V with a frequency in the range from 0.1 to 30 Hz, the typical results obtained are shown in Figure 18.6 [16]. Here, the vibration amplitude is that of the free end of a PLLA fiber with a diameter of 10 mm. The amplitude increased with decreasing applied AC voltage frequency and reached 20 mm at 0.05 Hz. In the case of DC voltage application, the strength of the applied DC voltage was varied from 0 to 800 V. Figure 18.7 shows the induced displacement generated by applying a voltage to PLLA fibers with a diameter of 10 mm [9]. In the case of the PLLA fiber with a diameter of 10 mm, when the applied voltage was 500 V, the displacement was greater than 30 mm; it increased with increasing applied DC voltage and reached 50 mm at 800 V. Moreover, a linear relationship was found between applied DC voltage and displacement in the range 0 V–800 V. On the basis of the experimental results, it was concluded that the shear strain locally induced by the electric field caused the bending of the entire PLLA fiber, and the fiber was observed to vibrate following the application of AC and DC voltages.

Catheters for Thrombosis Sample Exfoliation in Blood Vessels

363

Vibration Amplitude (µ m)

40

30

20

10

0 –1.0

–0.5

0.5 0.0 log f (Hz)

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1.5

Figure 18.6 Tip displacement (bending motion) against frequency of AC voltage applied to PLLA fiber.

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40 20 0 –20 –40 –60 –800

400 –400 0 Applied voltage (V)

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Figure 18.7 Tip displacement (bending motion) plotted against amplitude of DC voltage applied to PLLA fiber.

18.5 Elementary Demonstration of Prototype System for Catheters Using Piezoelectric Polymer Fiber To demonstrate the possible application of a piezoelectric polymer fiber, such as PLLA fiber, simple tools were developed. These were tweezers (for holding and extracting minute samples) and a catheter (for detecting, holding and extracting objects). Both make use of the piezoelectricity of the fiber. Electrically controlled tweezers using a pair of PLLA

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Figure 18.8 Illustration of PLLA fiber catheter and tweezers.

fibers (PLLA fiber tweezers) and a catheter using a PLLA fiber (PLLA fiber catheter) were designed (Figure 18.8) [9]. The concept is as follows. When the PLLA fiber touches an object, dynamic stress is generated because of the piezoelectricity. This stress induces displacement and charge on an electrode on the PLLA fiber. The induced charge becomes the signal for a subsequent electric field to be applied to the PLLA fiber. This field generates additional strain due to the bending motion of the entire PLLA fiber. Thus, the PLLA fiber can hold and extract the object with precision. In other words, the object can be held with a force that does not destroy it. 18.5.1

Preliminary Demonstration

Firstly, the trigger signal that needed to be produced when the PLLA fiber touches an object was examined. A urethane gel bead with a diameter of 300 mm was placed in a vessel. Then, an attempt was made to touch the object with PLLA fiber tweezers vibrated by an excitation vessel. Next, an attempt was made to grasp the bead by applying AC and DC voltages to the tweezers. A video recording of the test was made with a digital camera. Two still images are shown in Figure 18.9 [9, 15]. One shows the moment when the bead was released and the other shows the moment when the bead was grasped. Also, a detection signal due to piezoelectricity was generated when the tip of the fiber touched the object, as shown in Figure 18.9 [9, 15]. 18.5.2

More Realistic Model for Application of Piezoelectric Polymer Fiber to Catheter

From the results of the experiments described above, it was decided to consider a more realistic scenario. Here, it was assumed that a thrombosis had occurred in a blood vessel, as shown in Figure 18.10 [8, 9, 16]. The tip of a catheter was inserted at the point where a vessel branches off, and cells were exfoliated from the affected area for examination. For this demonstration, a PLLA fiber catheter was used (Figure 18.8) [9], operated according to the following procedure. Firstly, the PLLA fiber catheter was inserted under microscopy

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Figure 18.9 Detection on the basis of signal generated upon touching object with PLLA fiber.

Figure 18.10 Image of exfoliation of thrombosis from affected area using catheter.

[16]. A detection signal was generated when the tip of the fiber touched the object, which acted as a trigger for applying an electric field to the PLLA fiber catheter [9]. As a result, the fiber bent because of piezoelectricity. When the fiber came in contact with the bead, the signal induced by the PLLA fibers piezoelectricity was detected by a charge amplifier and digitized using a logic circuit. It was possible to detect the signal clearly [9]. It was also confirmed that this signal can be used as a trigger signal. An example is illustrated in Figure 18.11 [16], which shows the insertion of the fiber and the exfoliation and removal of a sample using the PLLA fiber catheter controlled by applying a voltage [8, 9, 16].

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Figure 18.11 Photographs showing exfoliation of sample using PLLA catheter controlled by applying voltage: insertion, exfoliation and removal of sample.

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A minute sample was successfully picked up, held and extracted. In this case, the PLLA fiber catheter was controlled by applying not only an AC voltage but also a DC voltage. The complicated motion of the PLLA fiber catheter involving vibration and rotation can be controlled by applying an AC voltage. Therefore, high precision was achieved in the manipulation of the electrically controlled PLLA fiber catheter to perform the very delicate operation of excising the affected area [8, 9, 16].

18.6 Summary The PLLA fiber obtained was driven by a piezoelectric effect when a voltage was applied to it; as a result, the entire PLLA fiber bent. We designed simple tools, such as a catheter and tweezers using PLLA fibers, that can be controlled by applying an applied voltage. Using the designed PLLA fiber tweezers and catheter, we successfully picked up, held and extracted minute samples. On the basis of these experimental results, it appears that it is highly possible to realize various applications using the piezoelectricity of chiral polymers such as the PLLA fiber.

References 1. Fukada, E. (2000) History and Recent Progress in Piezoelectric Polymers, IEEE Transactions of Ultrasonics, Ferroelectrics and Frequency Control, 47, 1110–9. 2. Fukada, E. (2006) Recent Developments of Polar Piezoelectric Polymers’’, IEEE Transactions on Dielectrics and Electrical Insulation, 13, 1110–9. 3. Wang, T. Herbert, J. and Glass, A. (1988) The Applications of Ferroelectric Polymers, Blackie, Glasgow. 4. Nalwa, H. (Ed.), (1995) Ferroelectric Polymers, Marcel Dekker, Inc., New York. 5. Lang, S. B. (2005) Guide to the Literature of Piezoelectricity and Pyroelectricity, Ferroelectrics, 321, 91–204. 6. Tajitsu, Y. (2002) Optical Rotatory Power and Light Modulation by Polylactic Acid Film, Mat. Res. Soc. Symp. Proc. Book, 698, 125–36. 7. Galetti, P., DeRossi, D. and DeReggi, A. (1988) Medical Applications of Piezoelectric Polymers, Gordon and Breach, New York. 8. Tajitsu, Y. (2006) Development of Electric Control Catheter and Tweezers for Thrombosis Sample in Blood Vessels Using Piezoelectric Polymeric Fibers, Polym. Adv. Tech., 17, 907–13. 9. Tajitsu, Y. (2008) Piezoelectricity of Chiral Polymeric Fiber and its Application in Biomedical Engineering, IEEE Transactions of Ultrasonics, Ferroelectrics and Frequency Control, 55, 1000–8. 10. Nye, J. (1985) Physical Properties of Crystals, Clarendon Press, Oxford. 11. Kobayashi, J., Asahi, T., Fukada, E., and Shikinami, Y. (1995) Structural and Optical Properties of Polylactic Acid, J. App. Phys., 77, 2957–72. 12. Aleman, C. Lotz, B. and Puiggali, J. (2001) Crystal Structure of the Alpha-form of Poly (l-lactide), Macromolecules, 34, 4795–801.

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13. Chawla, K. (1998) Fibrous Materials, Cambridge University Press, Cambridge. 14. Honda, M., Hayashi, K., Morii, K., et al., (2007) Piezoelectricity of Chiral Polymeric Fibers, Jpn. J. App. Phys., 46, 7122–4. 15. Tajitsu, Y., Kawai, S., Kanasaki, M., et al., (2004) Micro-actuators with Piezoelectric Polylactic Acid Fibers, Ferroelectrics, 304, 195–200. 16. Tajitsu, Y., Kanasaki, M., Tsukiji, M., et al., (2005) Novel Tweezers for Biological Cells Using Piezoelectric Polylactic Acid Fibers, Ferroelectrics, 320, 133–9.

19 Piezoelectric Poly(Vinylidene) Fluoride (PVDF) in Biomedical Ultrasound Exposimetry Gerald R. Harris Food and Drug Administration, Center for Devices and Radiological Health, USA

19.1 Introduction Since the discovery in 1969 of strong piezoelectricity in the fluoropolymer poly(vinylidene fluoride) (PVDF) [1, 2] this material has found use in an array of applications, including audio and ultrasonic transducers, accelerometers and vibration and touch sensors. One area in which piezoelectric PVDF has had a major impact is characterizing the acoustic output produced by biomedical ultrasound devices used in diagnostic and therapeutic applications. The ability to make accurate measurements of ultrasound exposure levels is of prime importance for assessing patient safety, improving imaging quality and increasing treatment effectiveness. The unique acoustic and mechanical properties of PVDF, including its availability in thin, flexible sheets, the ability to be spot poled and the similarity in acoustic impedance between PVDF and water, have contributed to its success in this application. In this chapter the development of PVDF transducers for characterizing medical ultrasound fields is described. Some of the material presented here is based on a previous effort of similar scope [3] but updated to include more recent work. The initial efforts in PVDF hydrophone development in the 1970s were motivated by the lack of a broadband sensor to measure the pulsed pressure fields produced by diagnostic ultrasound devices. The piezoelectric ceramic hydrophones available at the time were suitable for characterizing the continuous-wave and narrow band tone-burst pressures

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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produced by ultrasound physiotherapy devices. However, hydrophone housing reflections and multi-mode responses made them unacceptable for faithfully reproducing broadband diagnostic pulses that could have durations of only a few microseconds and bandwidths of several megahertz. The appearance of PVDF promised to overcome these problems and there were two basic designs of PVDF hydrophones that emerged: the needle, conceptually the same as existing ceramic hydrophones, and the spot-poled membrane. These two hydrophone designs are described in the next two sections. The following two sections present examples of experimental exposimetry work based on PVDF. These developments are discussed in two ultrasound application categories: diagnostic and therapeutic, the latter including acoustic surgery. This categorization is convenient for discussion purposes, but obviously some crossover exists.

19.2 Needle Hydrophone Design The PVDF needle hydrophone design was based on the similarly constructed piezoelectric ceramic hydrophone, an example of which is shown in Figure 19.1. The polymer implementation consists of a circular piece of thin (25 mm or less), metal-coated PVDF film, about 0.05 to 1.0 mm in diameter, attached to an insulating layer at the end of a hollowwalled, metal tube. The backing material behind the element typically has much higher acoustic impedance and absorption than water. Because the physical dimensions of the needle are very close to the dimensions of the active sensor element, the disturbance of the acoustic field is minimized. Therefore, this construction is particularly useful for spatial field plotting in the near field and for continuous wave measurements. Another advantage of the needle geometry is that it is readily adaptable for measurements in confined spaces such as in vivo. Other areas in which the needle hydrophone has found application are for miniature high frequency transducers and for the interrogation of transmitting arrays, element by element. A possible drawback of the needle design in some measurement situations is related to a roll-off in low frequency response due to diffraction at the needle tip [4–6]. Also, an anomaly in the frequency response due to an induced surface wave on the backing material has been observed [7]. Some work has transpired to reduce these effects, primarily via a modification of the housing [8, 9].

Figure 19.1 In this example of a needle hydrophone, the outside diameter of the hollow tubing to which the sensitive element is attached is 1.6 mm. The labeled scale dimensions are in centimeters.

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19.3 Spot Poled Membrane Hydrophone Design The advantageous properties of piezoelectric PVDF, especially the availability of unpolarized PVDF film in large, thin (approximately 5–25 mm), flexible sheets that could be spot poled, made a supported membrane approach a natural choice for designing hydrophones to measure millimeter-sized beams at ultrasonic frequencies above 1 MHz. Spot poling refers to the process of rendering only a small, central portion of the PVDF film piezoelectrically active. The dimensions of the annular framework holding the polymer sheet taut, usually 5–10 cm in diameter, can easily be made larger than the medical ultrasound fields being probed; thus the intrinsic broadband response of the polymer film can be exploited without introducing the complications associated with piezoelectric ceramic elements as mentioned previously. Arrays of sensitive elements have been fabricated on the membrane, but at present hydrophones with a single spot poled element are most often used. The sensitive element diameter of current hydrophones ranges from about 0.1 to 1 mm. Electrodes and electrical leads are vacuum deposited on the film. For 25 mm film the thickness resonance is approximately 40 MHz. An example device is pictured in Figure 19.2.

Figure 19.2 In this example of a spot poled membrane hydrophone, unpoled PVDF film 25 mm in thickness is mounted on a metal annular frame. Electrical leads on the upper and lower film surfaces adjoin overlapping 0.3 mm diameter electrodes. An electrical poling field applied to the leads renders the region of the film between the electrodes piezoelectrically active. An encapsulated preamplifier is located within the frame between the two coaxial connectors, one for preamp power in and one for signal out. The labeled scale dimensions are in centimeters.

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An advantage of the membrane design over the needle design is the ability to construct hydrophones with a wider, more uniform frequency response. One disadvantage of the basic single-layer membrane hydrophone design is that the electrical leads that lie on the polymer film surface are unshielded, which makes the hydrophone susceptible to RF interference. Also, the exposed leads can be a source of signal loss due to capacitive loading. These problems can be corrected by using a bilaminar membrane construction, in which two polymer sheets are first spot poled and then bonded such that the electrodes on the inner surfaces are aligned and the outer surfaces are coated with conductive material and connected to the signal ground to form parallel grounded planes. The main drawback of the bilaminar design, other than increased complexity of construction, is that because of the increased total film thickness, the thickness–resonance frequency and thus the useable bandwidth are reduced by a factor of two. However, the advantage of enhanced immunity to RF interference has made some form of this two-layer configuration the most extensively used. It has been demonstrated that PVDF based membrane hydrophones have negligible degradation in sensitivity with time, as well as good temperature stability and manufacturing reproducibility [10]. Also, a comprehensive model of the membrane hydrophone sensitivity transfer function has been developed [11]. This model incorporates multiple PVDF layers and amplifier characteristics, as well as the electrical cable and load.

19.4 Application to Diagnostic Ultrasound The early development of PVDF spot poled membrane hydrophones in the 1970s and early 1980s grew out of collaborative efforts between the Food and Drug Administration and the National Bureau of Standards (now the National Institute of Standards and Technology) in the United States [12–15], and also between the National Physical Laboratory (NPL) and GEC-Marconi Research Centre in the United Kingdom [16–19]. Both groups were motivated by the need for improving the acoustic characterization of diagnostic ultrasound devices. Their respective efforts [3] have helped to establish the spot poled membrane design as the de facto reference measuring hydrophone for medical ultrasonic fields. However, at this time other polymer hydrophone geometries were being explored with the same aim, the needle type being an important example [20–23]. Also, in a variation of the needle hydrophone design, the polymer film has been mounted in a baffle approximately 1 cm in diameter, with the centrally-located active diameter being closer to 1 mm in size [24, 25]. Thus the tube or similar structure to which the active polymer element was attached was significantly larger than the film dimensions. The baffle structure and attached housing were electrically shielded, and the housing contained an integral preamplifier. In other developments, voided PVDF film, that is film containing stretching-induced microvoids, has been used to construct hydrophones that provided an almost exact acoustic impedance match to water, as opposed to the factor of 2.7 greater for nonvoided PVDF [26]. Also, copolymers of vinylidene fluoride (VDF) have been used to construct needle [27], modified needle [9, 28] and spot poled membrane [29] hydrophones, as well as line hydrophones and arrays [30]. The copolymers have greater crystalline content than the homopolymer and thus greater sensitivity; however, the films are also more brittle and thus

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not as easy to work with in the membrane design [31]. The membrane hydrophone described in [29] was optimized for bandwidth and spatial resolution by using 4 mm thick trifluoroethylene copolymer film, P(VDF–TrFE), and a deposited 37 mm diameter spot poled electrode. From rise time measurements its bandwidth was estimated to be about 150 MHz. The ability of PVDF spot poled membrane hydrophones to make measurements over such a wide frequency range was recently studied more rigorously in [32], where a useable bandwidth of 140 MHz was reported for a hydrophone having a PVDF film thickness of 9 mm. To facilitate measurement of the acoustic output of ultrasound imaging devices in scanning modes, both one-dimensional (1-D) and two-dimensional (2-D) arrays have been developed [33–35]. The NPL 1-D array system described in [33] contained a PVDF membrane hydrophone with 21 elements of 0.4 mm diameter and 0.6 mm spacing. An upgrade from the design in [33] included a reduction in noise equivalent pressure to 1 kPa and a new 100 MHz A-D converter [34]. A 2-D array has been developed that uses a vinylidene fluoride copolymer membrane with an 8  8 pattern of elements having a diameter of 0.2 mm and a spacing of 1 mm [35]. Other hydrophone array designs have been built and used also, including annular [13, 36] and linear [30, 37] geometries. Further, optoacoustic imaging based on the generation of laser-induced ultrasound is a new and promising medical imaging modality, and a 128 element PVDF receiving array has been used in one system to record the acoustic transients [38]. For practical reasons exposure measurements to characterize the acoustic output of diagnostic ultrasound devices typically are made in water. Although water is a convenient and reproducible measurement medium, it is sometimes difficult to relate these measurements to values that would occur in tissue because of nonlinear propagation effects coupled with the differences in ultrasound attenuation between water and tissue [39]. To overcome this problem, an exposure measurement system has been developed comprising a milk based tissue-mimicking (TM) liquid filling a flexible sack that has a bilaminar PVDF membrane hydrophone mounted at one end [40]. The membrane design integrated well with this TM approach and allowed both accurate and physically relevant measurements to be made. Previously it was mentioned that the needle hydrophone geometry was readily adaptable for in vivo measurements. Other PVDF based designs for quantitative in vivo pressure measurements have been developed also. For example, in [41, 42] a seven-element PVDF linear array was described that measures in vivo exposures intervaginally during an obstetric ultrasound examination. The 0.5 mm diameter PVDF elements were spaced 1.5 mm apart along a stainless steel tube. A 16-element linear PVDF hydrophone array also has been designed, constructed and tested [43]. Each element was 0.5  1.0 mm with 3 mm center-to-center spacing. Although developed primarily for providing acoustic feedback during ultrasound hyperthermia treatments, the probe could be used for exposimetry also. PVDF exhibits pyroelectric as well as piezoelectric properties [44], a feature that has made the material useful for infrared sensing in motion detectors and thermal cameras. This attribute has been used recently in a thermally based method to measure ultrasonic power [45]. A 52 mm thick film of PVDF 60 mm in diameter was used as the sensor, which was backed with a rubber based material that generated heat due to ultrasonic absorption. Powers up to 1 W over a frequency range of 1–3.5 MHz were measured in this preliminary study.

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PVDF also has been employed in the construction of broadband transmitters for the purpose of calibrating hydrophones [46, 47]. In [46] the wide bandwidth of the 25 mm film allowed a single transducer to be used for a rapid calibration in the 1–40 MHz range; in [47] the technique was modified to extend the calibration range to 100 MHz.

19.5 Application to Therapeutic Ultrasound PVDF has also been used successfully as an acoustic pressure sensor in exposure measurements of various therapeutic ultrasound devices, including extracorporeal shock wave therapy (ESWT) and high intensity focused ultrasound (HIFU) surgery. ESWT devices, first used in the early 1980s for the fragmentation of urinary tract calculi in a procedure known as extracorporeal shock wave lithotripsy, produce focused, large amplitude pressure pulses. As in the case of diagnostic pulse measurements, the hydrophone must have a wide, uniform frequency response (up to 100 MHz is desirable) to record rise times that can be less than 10 ns. A small active element also is important, although not as critical as for diagnostic measurements, 0.5–1 mm being sufficient in most cases. In the case of lithotripsy fields, three additional hydrophone characteristics are important: robustness, to withstand the intense fields at the focus, which can be accompanied by cavitation; electrical shielding, to suppress RF interference from the shock wave generator and to provide insulation from conductive media in which the measurements sometimes must be made; and a dynamic range large enough to measure peak compressional pressures of up to 100 MPa, approximately ten times greater than those encountered in diagnostic fields [48]. Studies of pressure fields generated by commercial lithotripsy devices [49–56] have demonstrated the benefits of PVDF sensors in terms of measured pulse fidelity over conventional quartz or tourmaline ballistics pressure transducers. These latter transducers, while robust, have resonant frequencies less than 1 MHz and element dimensions of several millimeters [57, 58]. A variation on the needle design that has proved useful for lithotripsy field measurements was constructed by dipping the rounded tip of a needle in molten PVDF, cooling to form a film of varying thickness at the tip, poling via corona discharge and then coating the outer surface with evaporated metal [59, 60]. One feature of this construction technique is that the directional response is much broader than that of a flat needle design having similar dimensions (about 0.5 mm). This hydrophone, used for the studies described in [51, 54], was said to withstand several thousand lithotripsy pulses, but the rise time was still somewhat limited, being about 80 ns. More recently, in an in vitro study employing controlled cavitation to enhance stone erosion, this needle hydrophone was used to measure reflected pressure pulses and cavitation signals 6 mm from the focus to avoid damage [61]. For the work described in [49, 50, 52], bilaminar membrane hydrophones having a total thickness of 50 mm (2  25 mm) and active diameters of 1 mm or less were used. These hydrophones provided generally faithful reproductions of the lithotripsy pressure pulses. However, three potential problems were identified. In [52] an artifact was observed in the hydrophone waveform due to reflections at the hydrophone housing. In most cases this delayed disturbance would not interfere with the analysis of the main pulse detected by the

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hydrophone. However, because of this artifact, the size of the membrane support structure determines how far away from the lithotripsy beam axis measurements can be made, especially if that structure is highly reflective [62]. Two other possible problems were discussed in [50]. First, the thickness resonance of 24 MHz has the potential of causing errors in the measurement of rise time and peak compressional pressure, a situation dealt with theoretically in [63]. However, a comparison check with an 18 mm thick bilaminar hydrophone showed no significant differences for a rise time of approximately 30 ns and peak pressure of 40 MPa. The second problem was the aforementioned issue of hydrophone damage in the form of pitting of the metal electrodes and leads, attributed to the action of cavitation in the water measurement medium. While the performance of the hydrophone was not affected during the course of the study, it was concluded that eventually the hydrophone would have failed. Regarding cavitation, it has been noted that measurement of the negative pressure portion of the lithotripsy pulse may be affected by bubble formation at the surface of the hydrophone [62, 64]. Several modifications to the basic spot poled membrane design have been developed to try to circumvent the damage problem. In one approach, unmetallized PVDF film was spot poled and mounted in an enclosed chamber with an acoustic window. The chamber was filled with either a dielectric fluid [65, 66] or low resistivity electrolyte [67, 68]. Electrodes were located in the chamber away from the film, close enough to collect the pressureinduced charge, via capacitive coupling in the dielectric fluid case, but far enough away to minimize damage from the shock waves. In a different approach, the goal has been to produce an inexpensive, disposable membrane [69–73]. In [70, 71] a device with a bandwidth exceeding 20 MHz and an effective diameter less than 1 mm was described. In this device the resistance of the metal electrodes and leads is monitored to indicate when the film is about to fail and should be replaced. Several examples of the successful application of this hydrophone design have been reported, including its use in comparing rigid (i.e. conventional) and pressure-release ellipsoidal reflectors [62], in a study employing dual lithotripsy pulses to increase cavitation damage in vitro [74] and in characterizing a research lithotripter designed to reduce the potential for vascular injury [75]. To protect the hydrophone in [75] from cavitation damage, the PVDF membrane was encapsulated in castor oil and recalibrated. More recently a design based on [70, 71] was developed that was found to survive over 1000 lithotripsy pulses [73]. In an example of its use, changes in the spectrum of the scattered signal measured by the hydrophone were correlated with the degree of stone fragmentation [76]. In another low cost design [72], a commercial shock gauge element [77] was mounted on a housing containing degassed petroleum as a backing material for electrical isolation. While no single spot poled membrane design has yet emerged as the optimal choice, each of these variations has found use in lithotripsy field measurements. For in vivo measurements of lithotripsy pulses, both needle and membrane designs have been used. In one study, pressure waveforms in pigs were obtained using the needle probe described in [59, 60] and a commercial lithotripsy system [78]. Measurements were made at various axial distances and the results were compared to in vitro data. Two modifications of the single sheet, spot poled membrane design also have been reported. In one, previously metallized and spot poled PVDF film was formed

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into a hemispherical shell supported by a soft, low impedance backing [79]. The resultant bulb was mounted on the end of a stainless steel tube, with electrical connection being made via a coaxial cable as in the needle hydrophone. In another lithotripsy hydrophone design for measurements in pigs, a piece of metallized, spot poled film was stretched over a plastic ring having an outside diameter of 21.5 mm [80]. After connecting a coaxial cable, this miniature membrane hydrophone was then coated with silicone rubber before implantation to provide electrical insulation. Although lithotripsy is the most common application of ESWT, other uses exist, and PVDF has played a role in evaluating their performance. For example, a PVDF hydrophone was used to measure the acoustic field of a pneumatically driven device employed in orthopedics to treat heel pain [81]. Turning to HIFU, this technology is being used for localized destruction of a targeted tissue volume, primary via thermal ablation, with minimal damage to the surrounding region. Applications now in clinical use or under investigation include treatment of uterine fibroids and tumors in the prostate, liver, kidney, breast, brain and bone. Peak pressures are in the tens of megapascals, and treatment durations are from a few to tens of seconds. PVDF hydrophones can be used to characterize HIFU pressure fields, such as in [82], in which a needle hydrophone was employed to measure the reduction in focal intensity due to blockage of the HIFU beam by gas bubbles. In this case the peak levels were smaller than would be encountered normally, which afforded some protection to the hydrophone. Otherwise, to minimize the possibility of hydrophone damage at the highest peak pressure amplitudes, a burst mode is recommended with low duty factor and short pulse durations (e.g. 10–20 cycles) [83]. Another method to avoid the hydrophone damage that can occur when probing the focal region of a HIFU beam at high output levels is to use a small scatterer placed in the focus [84–86]. The reflected beam of much lower pressure is received by a hydrophone whose focus is centered on the scattering tip. In [84] the scatterer is a tapered glass rod with a tip diameter of about 100 mm. The tip is fixed at the focus of a piezoelectric ceramic receiver with a center frequency of 5 MHz and a radius of curvature of 7.5 cm; however, a 5 cm diameter, 10–15 cm focal length PVDF receiver is also being studied. In a similar approach, the scattering tip is placed at the center of a segmented, truncated spherical PVDF receiver with a 10 cm radius [85, 86]. The scattering tip is a fused silica optical fiber with a polyamide coating for protection. The total tip diameter is 83 mm. The receiving elements are poled 25 mm PVDF films. Each PVDF segment can be pulsed so the pulse-echo response can be used to aid in alignment of the fiber tip at the center of the sphere. In another HIFU application, short, high intensity pulses have been used for the cavitation mediated, nonthermal erosion of soft tissue [87, 88]. A bilaminar PVDF hydrophone was used to characterize the intense pressure fields. For this application the likelihood of thermal damage to the hydrophone is reduced because of the low duty factor and short tone bursts used (<25 cycles). However, damage from cavitation or other mechanical action is an issue. PVDF receivers are also being investigated for use during the clinical ESWT or HIFU procedure. In one in vitro lithotripsy study to measure the scattered signal from the targeted stone, it was shown that the acoustic signature of the scattered wave can be used to monitor the progress of stone fragmentation [89]. In another such lithotripsy

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study, both in vitro and clinical results indicated that a PVDF cavitation sensor showed promise for on-line monitoring of the degree of fragmentation of the stone [90]. In this latter study, the concept for the in vitro sensor was based on a design described in [91, 92], in which a hollow, open-ended cylindrical geometry, made possible by the flexible nature of PVDF film, was implemented for mapping the cavitation activity in devices such as ultrasonic cleaners. In a clinical HIFU application, a PVDF sensor was placed near the radiating surface of a spherically focused transducer [93]. In this low pressure region of the field the sensor was less susceptible to damage and its output was found to correlate with the power output from the transducer, thus providing a real-time monitoring feature. In another approach for monitoring the output power of a HIFU transducer during treatment, the pyroelectric property of PVDF has been used [94]. An 8 cm diameter bilaminar membrane having a total thickness of 120 mm was mounted in an annular frame and placed between the transducer radiating surface and the focus, tilted at a slight angle to avoid standing waves. The small amount of ultrasonic power absorbed in the membrane resulted in a temperature rise which was measured via both voltage and charge amplification. In this preliminary study the sensor produced promising results at a frequency of 1.7 MHz and ultrasonic powers up to 125 W. In a further study relevant to acoustic detection during clinical treatment, a broadband (0.1–10 MHz) PVDF strip sensor was placed in a sample of ex vivo bovine liver tissue to investigate the possibility of detecting cavitation signals from HIFU during clinical treatments, with the idea of exploiting cavitation activity to enhance lesion formation from HIFU [95]. It should be noted that both ESWT and HIFU measurements are still a significant challenge, and the ideal detector technology has yet to be found, primarily, again, because of the damage problem. Other approaches not relying on piezoelectric polymers that have been reported include spot poled ceramic [96], capacitive [97] and electromagnetic [98, 99]. Also, numerous uses of fiber optic probe hydrophones have been reported [64, 83, 100–109]. Of these device types the latter is said to give a more accurate measurement of the peak negative pressure because of the greater cavitation threshold at the water/fiber boundary compared to that of the polymer film and deposited metal. Also, a modified needle hydrophone has been introduced that is designed specifically to reduce the risk of cavitation damage in HIFU measurements by encasing the piezoelectric ceramic sensing element in a thin metallic coating (70 mm less) [110–112].

19.6 Conclusion The discovery of strong piezoelectricity in PVDF has led to the development of numerous polymer based acoustic sensors, in particular the spot poled membrane hydrophone design in its various embodiments. These hydrophones have become the primary devices for characterizing biomedical ultrasound fields. Their development has made possible the quantification of medical ultrasound exposure levels, and so enabled standards organizations to adopt a rational and systematic approach to address the problems of field measurements and device safety. Significant measurement challenges remain, particularly in the area of therapeutic ultrasound, but it is clear that PVDF will continue to play a pivotal role in the characterization of medical ultrasound devices.

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65. Granz, B. (1989) PVDF hydrophone for the measurement of shock waves, IEEE Trans. Electrical Insulation, 24, 499–502. 66. Granz, B., Holzapfel, R. and G. Ko¨hler, (1989) Measurement of shock waves in the focus of a lithotripter, Proceedings of the IEEE Ultrasonics Symposium, Montreal, Canada, 3–6 October 1989, 991–994. 67. Cathignol, D. (1990) PVDF hydrophone with liquid electrodes for shock wave measurements, Proceedings of the IEEE Ultrasonics Symposium, Honolulu, HI, 4–7 December 1990, 341–4. 68. Inose, N. and Ide, M. (1992) A miniature hydrophone for high acoustic pressures, Jpn. J. App. Phys., 31, 272–273. 69. Everbach, E.C. (1990) An inexpensive wide-bandwidth hydrophone for lithotripsy research, J. Acoust. Soc. Am., 87, suppl. 1, S128. 70. Schafer, M.E. (1993) Cost-effective shock wave hydrophones, J. Stone Disease, 5, 73–6. 71. Schafer, M., Kraynak, T. and Krakhman, V. (1994) Development of a cost-effective shock wave hydrophone, Proceedings of the IEEE Ultrasonics Symposium, Cannes, France, 1–4 November 1990, 1805–8. 72. Tavakkoli, J., Birer, A. and Cathignol, D. (1996) Development of a PVDF low-cost shock-wave hydrophone, Shock Waves, 5, 369–74. 73. Maxwell, A.D., Sapozhnikov, O.A. and Bailey, M.R. (2006) A new PVDF membrane hydrophone for measurement of medical shock waves, Proceedings of the IEEE Ultrasonics Symposium, Vancouver, Canada, 3–6 October 2006, 1608–11. 74. Sokolov, D.L., Bailey, M.R. and Crum, L.A. (2001) Use of two pulses to localize and intensify cavitation in lithotripsy. J. Acoust. Soc. Am., 110, 1685–95. 75. Zhong, P. and Zhou, Y. (2001) Suppression of large intraluminal bubble expansion in shock wave lithotripsy without compromising stone comminution: Methodology and in vitro experiments, J. Acoust. Soc. Am., 110, 3283–91. 76. Owen, N.R., Bailey, M.R., Crum, L.A. and Sapozhnikov, O.A. (2007) Identification of kidney stone fragmentation in shock wave lithotripsy, 3–6, New York, NY, 28–31 October 2007, 323–6. 77. Bauer, F. (1983) PVF2 polymers: Ferroelectric polarization and piezoelectric properties under dynamic pressure and shock wave action, Ferroelectrics, 49, 231–40. 78. Vergunst, H., Terpstra, O.T., Schroder, F.H. and Matura, E. (1990) In vivo assessment of shockwave pressures. Implication for biliary lithotripsy, Gastroenterology, 99, 1467–74. 79. Schafer, M.E., Kraynak, T.L. and Lewin, P.A. (1990) Design of a miniature in-vivo shock wave hydrophone, Proceedings of the IEEE Ultrasonics Symposium, Honolulu, HI, 4–7 December 1990, 1623–6. 80. Cleveland, R.O., Lifshitz, D.A., Connors, B.A., et al. (1998) In vivo pressure measurements of lithotripsy shock waves in pigs, Ultrasound Med. Biol., 24, 293–306. 81. Chitnis, P.V. and Cleveland, R.O. Acoustic and cavitation fields of shock waves therapy devices, in Therapeutic Ultrasound: 5th International Symposium on Therapeutic Ultrasound (eds Clement, G.T., McDannold, N.J. and Hynynen K.), AIP Conf. Proc. 829, Am. Inst. Physics, New York, 2006, 440–4. 82. Hosseini, S.H.R., Zheng, X., Vaezy, S. (1991) Experimental investigation of the effects of gas pockets on HIFU field, Proceedings of the IEEE Ultrasonics Symposium, New York, NY, 28–31 October 2007, 1309–12. 83. Zhou, Y., Zhai, L., Simmons, R. and Zhong, P. (2006) Measurement of high intensity focused ultrasound fields by a fiber optic probe hydrophone, J. Acoust. Soc Am., 120, 676–85. 84. Kaczkowski, P., Cunitz, B., Khokhlova, V. and Sapozhnikov, O. (2003) High resolution mapping of nonlinear MHz ultrasonic fields using a scanned scatterer, Proceedings of the IEEE Ultrasonics Symposium, Honolulu, HI, 5-8 October 2003, 982–5.

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Section V Dielectric Elastomers

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

20 Dielectric Elastomer Actuators: Fundamentals Roy Kornbluh, Richard Heydt and Ron Pelrine SRI International, USA

20.1 Introduction Originally developed as muscle-like actuators for small robots [1], dielectric elastomers are a type of electric-field-activated electroactive polymer with demonstrated performance that suggests that they may have wide applicability for use as actuators, generators and sensors. In particular, dielectric elastomer actuators have demonstrated both strain and specific energy densities far greater than those of any other field-activated material, including single crystal piezoelectrics [2]. This performance, combined with other potentially desirable features such as high compliance, easy integration with body tissue and the option to use low cost environmentally tolerant (and most likely biocompatible) materials, makes dielectric elastomers attractive for many medical applications. Dielectric elastomer technology was pioneered by Pelrine and others at SRI International [3]. SRI International (SRI) has been developing dielectric elastomer transducers since 1992. More recently, an increasingly large number of research groups around the world have been developing dielectric elastomer technology and devices that use this technology. For example, the Electroactive Polymer Actuators and Devices 2008 Conference (SPIE) featured more than 28 papers on dielectric elastomers by more than 20 research groups from around the world. In 2004, Artificial Muscle Inc. was spun off from SRI in order to commercialize the technology. Companies are also emerging that are exploiting the technology for specific applications.

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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This chapter introduces dielectric elastomer technology so that its application to medical devices, discussed in subsequent chapters, can be better understood.

20.2 Basic Principle of Operation Dielectric elastomer transducers are based on the electromechanical response of an elastomeric dielectric film with compliant electrodes on each surface. These transducers may be actuators, generators, or sensors. In all cases, the basic structure is the same. Actuators based on dielectric elastomer technology operate on the simple principle shown in Figure 20.1. When a voltage is applied across the compliant electrodes, the polymer shrinks in thickness and expands in area.

Figure 20.1 Principle of operation of dielectric elastomer actuators. (a) Functional element of dielectric elastomer actuators. Polymer film compresses in thickness and expands in area when a voltage is applied across the film. (b) Typical thickness or planar strain in response to applied electric field for a film with no external loads.

The polymers are largely incompressible. Therefore, the electrodes must be compliant to allow the film to strain in area as well as in thickness. The observed response of the film is caused primarily by the interaction between the electrostatic charges on the electrodes [4, 5]. Simply put, the opposite charges on the two electrodes attract each other, while the like charges on the electrodes repel each other. Based on this simple electrostatic model, it is possible to derive the effective pressure produced by the electrodes on the film as a function of the applied voltage [6]. This pressure, p, is: p ¼ er eo E2 ¼ er eo ðV=t Þ 2

ð20:1Þ

where er and eo are the permittivity of free space and the relative permittivity (dielectric constant) of the polymer, respectively; E is the applied electric field; V is the applied voltage; and t is the film thickness. The response of the polymer is functionally similar to that of electrostrictive polymers, in that the response is directly related to the square of the applied electric field. Electrostrictive polymers differ from dielectric elastomers in

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that there is an additional term contributing to the dielectric constant in Equation (20.1) (see, for example, Chapter 16 on the fundamentals of electrostrictive polymer actuators). While the equation for actuation stress is straightforward, the resultant strains produced in the polymer are dependent on the boundary conditions and loads on the polymer. Further, the strains depend on the elastic modulus of the polymer, which may be nonlinear for elastomers at large strains. A further complication is that the polymers are often actuated with large initial pre-strains. Such pre-strains may be anisotropic and thus may cause the effective elastic modulus to also be anisotropic. For these reasons it is not possible to write a simple general equation for the resultant strain. Several researchers have developed models that include the strain response of dielectric elastomers (e.g. Goulbourne et al. [7]). Dielectric elastomers also can be operated in generator mode: a bias voltage is applied across the polymer and the deformation of the polymer causes a change in the capacitance. If the capacitance decreases, there is a net production of electrical energy. For example, in large strain deformations the amount of electrical energy generated per unit volume of material, e, is: e ¼ er eo E2 ¼ er eo ðV=t Þ 2

ð20:2Þ

The operation of dielectric elastomers as generators is discussed in more detail by Pelrine et al. [8]. The same transducer structure can serve as either an actuator or a generator, depending on how it is electrically driven. Therefore, the relationship between motion and energy of deformation can be modulated. In other words, the transducer can be a variable impedance device in which stiffness and damping are electrically controllable. Since the same transducer structure can serve as an actuator, generator, sensor or variable impedance device, the transducer can be said to be multifunctional. The use of dielectric elastomers as ‘variable stiffness’ devices is discussed in detail elsewhere [9]. Since the electric impedance and the output of a dielectric elastomer transducer are related to the deformation of the polymer film, it follows that these transducers could also be used to detect strain. For example, a linear actuator could be designed such that film capacitance is directly related to the amount of actuator linear motion. Rosenthal [10] discusses the use of dielectric elastomers as a sensor.

20.3 Dielectric Elastomer Materials While the performance of a dielectric elastomer transducer depends upon the dielectric and elastomeric properties of the polymer material, a great many polymer materials can be used. Because of this flexibility, unlike most other electroactive polymers (EAPs), different polymer materials can be selected for different applications, depending on the desired performance and physical properties. The best performing materials (those with greatest strains) are based on commercially available formulations of silicone rubber (polydimethyl siloxane) and acrylic elastomers such as the VHB series from 3M Corp. (Minnesota, USA) [11].

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Figure 20.2 Planar strain of 300 % produced in an acrylic film. The photographs show a section of a long electrode (dark area).

Both silicone and acrylic materials have produced more than 100 % strain. A planar strain of 300 % is shown in Figure 20.2. The specific energy density associated with such large deformations is greater than that produced by any other field-induced actuator technology. Since the pressure response is due to the movement of electrons as manifested in electrostatic charges (electric-field activated), it is very fast. The strain response is limited in speed by the damping of the polymer and the modal frequencies that result from the elasticity of the polymer and the driven inertia. More details are given by Kornbluh et al. [12]. The performances of the best silicones and acrylics are summarized in Table 20.1. The performance of both the acrylic and silicone elastomers was shown to be enhanced by applying a ‘pre-strain’, or initial state of strain to the material [2].

Table 20.1

Performance and characteristics of the best silicone and acrylic elastomers

Parameter

Acrylics

Silicones

Comment

Maximum actuation strain (%)

380

120

Maximum actuation pressure (MPa) Maximum specific energy density in actuation (MJ/m3) Maximum frequency response (Hz)

7.2

3.0

Free strain. Strain in generator and sensing modes can be greater than that in actuation. Blocked stress.

3.4

0.75

>50 000

>50 000

Greatest energy density of all field-activated materials; can be greater in generator mode. Small-strain acoustic measurements; frequency response is very dependent on strain and size; full response of acrylic is generally much lower (less than 1 kHz), due to viscoelasticity (see ‘Mechanical loss factor’ below). (continued )

Dielectric Elastomer Actuators: Fundamentals Table 20.1

391

Continued

Parameter

Acrylics

Silicones

Comment

Maximum electric field (MV/m)

440

350

Relative dielectric constant Dielectric loss factor Average elastic modulus (MPa) Mechanical loss factor

4.8

2.5 3.0

0.005 0.5–3.0

<0.005 0.1–2.0

0.18

0.05

Maximum electromechanical coupling, k2 Maximum overall electric to mechanical transduction efficiency (%) Durability (cycles)

0.9

0.8

>80

>80

>10 000 000

>10 000 000

Maximum fields are realizable only in uniform films with few defects. Measured at 1 kHz; small drop-off at higher frequencies. Measured at 1 kHz. Elastic modulus is typically nonlinear with respect to strain. Loss factor of acrylics varies with strain and other conditions. Similar to that of the best fieldactivated materials such as singlecrystal piezoelectric ceramics. Assumes ideal driving electronics at ideal frequency (about 10 Hz); efficiency is frequency dependent. Durability is highly dependent on how close the driving voltage is to the maximum field. Short-term range; long-term range has yet to be determined.

Operating range (°C)

10 to 90

100 to 260

20.4 Transducer Designs and Configurations The basic element shown in Figure 20.1 can be incorporated into a wide variety of transducer configurations. For example, the film and electrodes can be formed into a tube, rolled into a scroll, stretched onto a frame or laminated to a flexible substrate to produce bending. Several examples of actuator configurations are shown in Figure 20.3. Which configuration is best depends on the application and properties of the film. In many cases these configurations are similar to those that are well known for piezoelectric ceramic devices. However, in many other cases, the configurations are quite different and exploit the unique properties of dielectric elastomers such as large strain, flexibility, coupling between different directions of motion and the need to hold a ‘prestrain’. For example, the bow tie and spider devices couple two in-plane directions of deformation to a single output. The framed transducers are well-suited to maintaining prestrain and incorporating multiple individually addressable active regions into a single flexible substrate. While the examples in Figure 20.3 show a wide variety of designs, there are many other possible designs and variations. The subsequent chapters of this book discuss in more detail how these and other designs may be incorporated into medical devices.

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Figure 20.3

Representative dielectric elastomer actuator configurations.

20.5 Operational Considerations Dielectric elastomer EAPs are highly capacitive electrical loads and typically operate at high voltages (300 V to 5 kV, depending on the film thickness). High voltage operation of a capacitive load is fast and efficient but can pose design challenges for biomedical and other applications. In small devices, the total electrical power and stored energy do not pose much of a safety risk, although in larger devices safety issues must be considered. It is not the voltage per se that poses the risk but rather the leakage current through the tissue. Proper packaging can keep this current well below levels that pose a risk under normal operation. In many non-EAP applications, packaging minimizes the risk associated with high voltage (CRTs and fluorescent lighting are examples of high voltage devices that are used extensively and safely). Similarly, integrating voltage converters with the dielectric

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elastomer transducer can mitigate some of the risk. In the future, such converters might be integrated with the transducer on the same polymer substrate as that which constitutes the actuator. Regardless of whether or not the voltage converter is integrated, most applications will require both voltage converters and high voltage switches (e.g. transistors) as part of the entire system. Suitable components are commercially available or expected to be in the near future. The cost is not significantly greater than that of other common electronic components. Applications that require more sophisticated control, high frequency operation or charge recovery for improved efficiency will require additional circuitry. A more detailed discussion is given by Kornbluh et al. [12].

References 1. Kornbluh, R., Pelrine, R. and Joseph, J. (1995) Elastomeric Dielectric Artificial Muscle Actuators for Small Robots, Proceedings of the Third IASTED International Conference on Robotics and Manufacturing, Cancun, Mexico, 14–16 June 1995, 1–6. 2. Pelrine, R., Kornbluh, R., Pei, Q. and Joseph, J. (2000) High-Speed Electrically Actuated Elastomers with Over 100 % Strain, Science, 287:5454, 836–9. 3. Pelrine, R. and Kornbluh, R. Introduction: History of dielectric elastomer actuators in Dielectric Elastomers as Electromechanical Transducers (eds Carpi, F., DeRossi, D., Kornbluh, et al.), Elsevier, Oxford, UK, 2008. 4. Pelrine, R., Kornbluh, R. and Kofod, G. (2000) High-Strain Actuator Materials Based on Dielectric Elastomers, Adv. Mat. 2000, 12:16, 1223–5. 5. Kornbluh, R., Pelrine, R., Joseph, J., et al. (1999) High-Field Electrostriction of Elastomeric Polymer Dielectrics for Actuation, Smart Structures and Materials 1999: Electroactive Polymer Actuators and Devices, Proceedings of SPIE (ed. Bar-Cohen, Y.), 3669, 149–61. 6. Pelrine, R., Kornbluh, R., Joseph, J. and Marlow, J. (1998) Analysis of the Electrostriction of Polymer Dielectrics with Compliant Electrodes as a Means of Actuation, Sensors and Actuators A: Physical, 64, 77–85. 7. Goulbourne, N., Mockensturm, E. and Frecker, M. (2005) A nonlinear model for dielectric elastomer membranes, J. Appl. Mech., 72, 899–906. 8. Pelrine, R., Kornbluh, R., Eckerle, J., et al. (2000) Dielectric Elastomers: Generator Mode Fundamentals and Applications, in Smart Structures and Materials 2001: Electroactive Polymer Actuators and Devices, Proceedings of SPIE (ed. Bar-Cohen, Y.), 4329, 148–56. 9. Pelrine, R. Variable stiffness mode: devices and applications, in Dielectric Elastomers as Electromechanical Transducers (eds Carpi F., DeRossi D., Kornbluh R., et al.), Elsevier, Oxford, UK, 2008. 10. Marcus Rosenthal, Neville Bonwit, Charlie Duncheon and Jon Heim (2007) Applications of dielectric elastomer EPAM sensors, Proceedings of SPIE, 6524, 65241F. 11. Kornbluh, R., Pelrine, R., Pei, Q., Oh, S. and Joseph, J. (2000) Ultrahigh Strain Response of Field-Actuated Elastomeric Polymers in Smart Structures and Materials 2000: Electroactive Polymer Actuators and Devices (EAPAD), Proceedings of SPIE (ed. Bar-Cohen, Y.), 3987, 51–64. 12. Kornbluh, R., Pelrine, R., Pei, Q., et al. Application of Dielectric Elastomer EAP Actuators, in Electroactive polymer (EAP) actuators as artificial muscles: reality, potential and challenges, 2nd edn., (ed. Bar-Cohen, Y.), SPIE Press, Bellingham, Washington, 2004.

21 Biomedical Applications of Dielectric Elastomer Actuators John S. Bashkin1, Roy Kornbluh2, Harsha Prahlad2 and Annjoe Wong-Foy2 2

1 Fremont, CA, USA SRI International, USA

21.1 Introduction Dielectric elastomer actuators are also known as ‘Electroactive Polymer Artificial Muscle’ or EPAM*, for short. As the name implies, this technology can replicate some of the form and function of natural muscle, and do it under electrical control. Thus, there would seem to be great potential for this new transducer technology for biomedical applications. In this chapter, we describe several recent efforts to apply EPAM technology to biomedical applications that we have undertaken in partnership with several collaborators, so illustrating the breadth of potential applications and actuator configurations. This chapter is not intended to be an exhaustive compilation of all biomedical devices that could benefit from EPAM. Rather, the focus is on specific case studies with which the authors have first hand experience. These case studies serve to show the great promise, versatility and challenges of the use of EPAM for biomedical devices. In each case, the unique capabilities offered by EPAM, as compared to traditional transducer technologies as well as other electroactive polymers, are highlighted. There are, however, many other potential uses for EPAM technology as sensors for blood pressure monitoring, constricting bands for compression *

EPAMTM is a trademark of Artificial Muscle, Inc.

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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therapy and blood pressure cuffs, and for use in pumps and compressors for, for example continuous positive airway pressure (CPAP) and other respiratory therapy devices. To start the focus is on an application close to a commercial product, that of a diaphragm pump that can be used for external or implantable medical device pumps. As part of the pump case study, discussed in some detail are the EPAM actuator configuration, known as the Universal Muscle ActuatorTM (UMA). The UMA is a versatile transducer configuration that was developed by Artificial Muscle, Inc., the company formed by SRI International to commercialize EPAM technology. While the discussion is on the pump application, many other applications of UMA cartridges are possible, as they represent a general platform for linear actuation. Our focus on this type of actuator also shows the design rules and performance characterization that is needed if EPAM is to be used by biomedical device designers. Next, devices based on thickness-mode actuation are presented. Thicknessmode actuation has been useful for haptic feedback for which there are several potential medical device applications. These completely soft actuators can only be made with polymer materials and thus are an example of something that cannot be made with conventional technology. Other EPAM film based actuators for implantable applications are then described, including an artificial diaphragm and an artificial facial muscle for eyelid control. These latter applications are farther from commercial products but represent initial attempts to realize the potential of EPAM to truly act as artificial muscles and restore lost function due to injury or disease. To conclude there is a description of how EPAM can be used to imitate in vivo conditions in an in vitro cellular assay device. In general, the applications highlight the unique combination of large strain, fast response and high efficiency operation that is available with dielectric elastomer transducers. While not discussed in detail in this chapter, it is also noted that many of the actuator configurations presented here could be run in reverse as sensors or power generators for biomedical applications (such as measuring and harvesting human motion and other physiological activities).

21.2 UMA Based Actuators and Their Application to Pumps As described more fully elsewhere in this volume, the basic operation of EPAM consists of a thin film of a dielectric elastomer, typically a silicone or acrylic or other low-durometer material, that is coated on both faces with a compliant electrode layer. The resulting construct is a stretchable capacitor and the application of an electric voltage across the dielectric elastomer creates a voltage gradient between the electrodes. The applied voltage and thinness of the dielectric layer result in a large electrostatic potential force between the electrodes. As opposite charges attract and like charges repel in reponse to the applied electric field (Maxwell pressure), the elastomer film is squeezed thinner and expands in area [1]. The conversion of the electrostatic potential energy into mechanical work is efficient, allowing the basic phenomenon to be harnessed into useful actuator designs. To facilitate and standardize the EPAM actuator design process, Artificial Muscle, Inc. has developed a standard, modular building block for linear acutators that can be deployed for many applications with minimal customization. The Universal Muscle Actuator

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397

(UMA) shown in Figure 21.1 consists of a planar ring of EPAM held in a rigid frame and then assembled into a final actuator. The basic film cartridges are manufactured in a range of diameters and with varying numbers of elastomer film layers. A larger diameter gives a longer maximum actuator stroke while varying the number of layers within a planar cartridge increases the maximum restrained force available. In addition, multiple cartridges can then be stacked in parallel to further increase the actuator force, or used in a two-phase configuration to also increase stroke. By using standardized configurations, the same manufacturing process can be used to create a wide array of linear actuators as replacements for traditional solenoids or electric motors. A dual phase actuator design is shown in Figure 21.1, with two EPAM cartridges operating in opposition to each other in a two phase, pull–pull configuration, but single phase actuators, operating aganst a bias spring or other loading mechanism, are also possible. In fact, the examples in this section are based on devices that use a bias spring.

Figure 21.1 Basic construction of a Universal Muscle Acutator. Left: Assembly of a planar two-layer EPAM film cartridge; Middle: Assembly of two biased cartridges into a two-phase actuator; Right: Assembled UMA actuator. (Reprinted with permssion from Artificial Muscle, Inc.).

In its standard form, the UMA consists of two independent stacks of film, attached at their centers and separated by a lightweight spacer. Each stack of film attaches to a frame that gives structure to the film and usually used to anchor the UMA to the host system. An output disc is also attached to each stack of film, through which force and stroke are transmitted from the expanding polymer film to the mechanical load. A stack of film sandwiched between two frames and two output discs comprises a cartridge. The UMA can directly drive linear positioning devices at frequencies from DC to several hundred Hertz and does not require conversion of high speed rotary motion, as is commonly required for conventional rotary motors. This ‘direct drive’ capability allows designers to eliminate complicated gear reduction and motion conversion power trains, often leading to reduction in parts count by up to 50 % over traditional designs. The standard UMA actuators exhibit maximum force at zero displacement (restrained) that then decreases linearly to zero at the maximum displacement (unrestrained), as shown in Figure 21.2. The first step in the design process is to map the force–stroke curve for a particular application against the data illustrated in Figure 21.2. This step helps define the optimum UMA diameter, number of EPAM layers and number of UMA cartridges required. The other variable not included in Figure 21.2 is the required operating lifetime for the UMA. In general, the lower the applied electric field, then the longer the UMA lifetime. Designers can optimize the UMA operating characteristics by balancing stroke, operating frequency and cycle life by choosing the apropriate UMA dimension, number of EPAM layers and applied voltage.

398

Biomedical Applications of Electroactive Polymer Actuators 2.5 kV

1.2

2.0 kV 1.0

Zero Stroke

1.5 kV Operating Mid-point

Force (N)

0.8

1.0 kV

0.6 0.4

Zero Force

0.2 0.0 0.0

0.2

0.4

0.6

0.8

1.0

Stroke (mm)

Figure 21.2 Typical force vs stroke profile at different voltages for UMA actuators made with a single D50 UMA cartridge (Reprinted with permssion from Artificial Muscle, Inc.).

Fluid pumps for drug infusion, inflation of bladders or other applications are some of the most commonly pursued medical device applications of EPAM. Desired product attributes such as improved power efficiency, noise-free and vibration-free operation and light weight are commonly stated goals of medical device companies that are pursuing EPAM based systems. The potential for improved power efficiency of an EPAM pump over traditional pump mechanisms translates, for example, into smaller battery requirements for a long term implantable application. This is an important system level consideration because the battery can be the single most expensive component in an implantable system. The fluid power required for medical pumps, and hence the force and stroke for the pump actuator, is typically at the higher end of what EPAM technology has been able to achieve to date, but continued improvement in the performance of EPAM technology is beginning to make these applications feasible. Most of the EPAM based pump designs to date have been for diaphragm pumps, however, and other volumetric pump designs, such as peristaltic or syringe pumps, are preferred for many therapeutic applications. To illustrate the design process, we will describe at a high level the design approach to a typical EPAM based pump, using the UMA cartridge as the basis for a diaphragm pump actuator. In defining the engineering requirements for a diaphragm pump, the discussion begins with the maximum operating pressure (P) and flow rate (Q) required. These values relate directly to the force (F) and stroke (S) required of the actuator through the pumping piston area (A): P ¼ F=A

ð20:1Þ

Q ¼ V f

ð20:2Þ

V ¼ AS

ð20:3Þ

Biomedical Applications of Dielectric Elastomer Actuators

399

where the frequency of the pump is f and the volume of each pump bolus is V. The pump piston diameters are typically in the range 2–12 mm. The other limitation in the pump design is the lifetime requirement, and the relationship between EPAM lifetime and the applied electrical field (E) used to actuate the UMA cartridges. This relationship is not linear but, as with all capacitors, shows power dependence with an exponent that depends on the dielectric elastomer used. Biomedical pumps will typically require lifetimes in the 10–1000M cycle range. This lifetime requirement therefore dictates the maximum electric field used be limited to around £50 MV/m (although current research is enabling this limitation to be exceeded). A given UMA size will, in turn, display a linear relationship between force and stroke at a given applied electrical field (Figure 21.2). The three requirements of electrical field, force and stroke will together determine the UMA size and whether multiple cartridges are preferred. The pump flow rate requirement along with the stroke available from the chosen UMA will determine the pumping frequency. Typically, biomedical pumps will require somewhere between 2 and 20 D25 cartridges (the designation ‘DXX’ indicates that the cartridge has an outer diameter of XX mm) operating between 1 and 250 Hz. Additional experimental optimization will be required to explore the impact of the bias spring, diaphragm stiffness and so on. In addition, the pump will operate most efficiently at resonance. For a D25 UMA pump, this resonance is generally in the 75–250 Hz range. For example, consider a pump that requires a blocked pressure of 0.007 N/mm2 (1 psi) and a free flow rate of 2 cc/min. A commercially available diaphragm pump head used to demonstrate feasibility has a pumping diaphragm radius of 2.1 mm. For a pressure of 0.007 N/ mm2, this requires a force on the diaphragm of 0.097 N. While the free stroke displacement of a D25 UMA cartridge is known, the displacement of the UMA diaphragm pump coupled system must be characterized experimentally. This displacement and the radius of the pumping chamber give the volume of each fluid bolus. The maximum operating voltage for a D25 UMA is 2.0 kV. Operation of the pump over a sweep of frequencies at this voltage while monitoring the free flow rate provides an initial indication of the required operating conditions. Figure 21.3 shows the operating characteristics of this pump at two different operating voltages and frequencies, demonstrating the feasibility of meeting the user requirements at approximately 15 Hz and 2.0 kV. Increasing the voltage raises the blocked pressure generated by the pump, and increasing the frequency increases the free flow. By adjusting both the operating voltage and the operating frequency, these pumps can provide a large dynamic range of flow rates with high accuracy. Further optimization involves characterizing the system fully as a function of frequency to identify system resonances. Operating at resonance will maximize the mechanical and electrical power efficiencies of the system. In addition, the lifetime of the pump can be substantially improved by dropping the operating voltage down from 2 kV. Increasing the number of UMA cartridges from one to two allows the same pump performance to be achieved at 1.75 kV, and from two to four cartridges drops the voltage further to 1.4 kV. The drop in voltage from the initial 2 kV to 1.4 kV translates into an increase in the cumulative work life of the pump by roughly two orders of magnitude. The final pump design has four UMA cartridges

400

Biomedical Applications of Electroactive Polymer Actuators

Figure 21.3 Typical pressure vs flow rate curve for a UMA based diaphragm pump (Reprinted with permssion from Artificial Muscle, Inc.).

operating at 1.4 kV maximum and 22 Hz. The power efficiency of the entire pump system was three to five times higher than other diaphragm pumps with similar operating characteristics.

21.3 Mechanical Stimulation Using Thickness-Mode Actuation A thin, compliant, conformal vibrating actuator is desirable in many biomedical applications. The thickness-mode actuation technique [2], a recent configuration of dielectric elastomer actuation, greatly simplifies mechanical stimulation devices. Using this technique, virtually any desired pattern of bumps and troughs can be created on a single substrate by simply patterning the electrodes on one surface of the dielectric elastomer. The entire structure can be attached to a rigid frame or laminated to a flexible foam backing. In this way, a thin flexible skin that covers large or small areas and contains finely patterned electrodes can be constructed. This skin can be quite rugged and has no ‘hard points’ because it is made entirely of rubbery materials. Although the device gives out-of-plane motion, it can be fabricated using only twodimensional (2-D) patterning. The result is a spatially distributed, thin, planar actuator or sensor composed entirely of soft, rubbery materials that can provide mechanical stimulation (or sensing). These characteristics make the thickness-mode configuration ideally suited for some of the biomedical applications described below. It should be noted that although the basic principles and applications are discussed here with respect to dielectric elastomers, many of the same techniques can be directly applied to other types of electroactive polymers (EAPs) and, in fact, even to some non-EAP smart

Biomedical Applications of Dielectric Elastomer Actuators

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materials. However, since many applications require relatively high frequencies and strains, we believe that dielectric elastomers are ideally suited to such applications due to their fast speed of response, environmental stability and efficiency (resulting in low heat build-up). In the primary configuration, the active dielectric elastomer polymer film is bonded or coated with a thicker passive layer, such that changes in the polymer thickness during actuation of the dielectric elastomer device are at least partially transferred to the passive layer. The passive layer in this case has no electric field applied to it, and all the active properties of the laminate structure come from the underlying dielectric elastomer layer. Nevertheless, the passive layer and the dielectric elastomer layer are mechanically coupled, so area and thickness changes in the dielectric elastomer layer induce shearing forces in the passive layer that change the thickness of that layer. Therefore, the change in thickness in the passive layer amplifies, in absolute terms, the displacement produced by an increase in area or a change in thickness of the dielectric elastomer layer. Note that the thicknessmode technique does not actually require any change in thickness of the dielectric elastomer (or other EAP) layer to produce a change in thickness in the passive layer. A schematic diagram of this type of device and the results of thicknessmode actuation are shown in Figure 21.4. In the figure, depressions indicate electroded regions and the raised lines indicate areas bordering active regions. The grid pattern arises from the electrode configuration on either side of the active polymer film. The pattern on the thickness-mode devices can be adapted in any given application. The electrodes can also be addressed spatially to create a ‘travelling wave’ along the surface of the entire device. These devices are typically operated under low to moderate actuation strains. They have excellent frequency response characteristics in the range most sensitive to human skin (below 100 Hz). An obvious application is as a massager in foot or lumbar massaging applications. In contrast to the currently marketed rigid electromagnetic massaging actuators, these actuators match better to the impedance of skin and feel more comfortable against the human body. Lightweight, portability and mechanical flexibility are also inherent assets. For example, a large, thin, compliant and lightweight elastomeric pad can be used as a portable massaging pad that can be draped over a chair or mattress. The material’s compliance and absence of hard points is also an advantage for massaging cuffs, braces and so on, such as those used in sportswear and for physiotherapy. Flexible and wearable massaging cuffs, gloves or armrests may also aid surgeons during long procedures. For example, we are investigating integration of vibratory massagers as the cushioning layer in handles or armrests during telesurgery for feedback of various forces during surgery. Another attractive feature of the thickness-mode dielectric elastomer is that mechanical loading (in a direction normal to the plane of the device) can result in even greater actuation pressures. The mechanical pre-load load serves to decrease any viscoelastic or other losses in the passive layer and in the transfer of the actuation pressures from the film more directly through the passive layer (at the cost of thickness change). This feature is relevant for footwear applications where human weight would serve as a pre-load to improve the vibrotactile forces felt through the thickness-mode dielectric elastomer.

402

Biomedical Applications of Electroactive Polymer Actuators

Figure 21.4 Thickness-mode actuation: (top) underlying electrode grid pattern and schematic of basic element; (bottom) resulting surface change of acrylic dielectric elastomer with electrodes in grid pattern (Reprinted with permission from Prahlad, H., Pelrine, R., Kornbluh, R. et al. Programmeable surface deformation: thickness-mode electroactive polymer actuators and their applications, Proceedings of SPIE, 5759, 102. Copyright (2005) SPIE).

An application for these active insoles that has been studied in some detail involves using them to enhance the sense of balance in elderly and diabetic patients. Clinical trials have shown that adding small amounts of vibratory noise to an elderly or diabetic person’s foot can act as a carrier to carry the sensation felt at the feet over their lower sensory

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403

threshold. In particular, vibrations induced at the feet have been shown to reduce the amplitude of posture sway of a person while standing, thereby reducing the likelihood of falls and injury [21.3]. EPAM thickness-mode actuation could provide the vibration stimulus for such applications by uniquely enabling a soft, comfortable insole that offers both cushioning and vibration functionality. Other applications of thickness mode actuation that have not been fully investigated to date involve shifting pressure points in bedridden patients to prevent bedsores, massaging functionality to prevent thrombosis, and in tactile feedback devices including refreshable Braille devices [2].

21.4 Implantable Artificial Diaphragm Muscle Recently, Bashkin, et al. reported [4] the initial development of an artificial implantable diaphragm envisioned for use in patients suffering from diaphragmatic and/or phrenic nerve failure due to traumatic spinal cord injury or one of several muscle wasting diseases. The concept is outlined in Figure 21.5 which shows (a) the implantable actuator powered through an external induction coil and (b) an illustration of the EPAM actuator prototype. The artifical diaphragm operated in an analagous fashion to the natural muscle. The EPAM film was held in a rigid frame under tension, which served to keep the film flat (i.e. ‘muscle contracted’) against the positive pressure bias from the abdomen to the chest cavity. This corresponded to the peak inhalation state. When voltage was applied to the EPAM, the film expanded in area and the positive abdominal pressure bias caused it to distend upwards (i.e. ‘muscle relaxed’). This state corresponded to the full exhalation state. When power was turned off, the film contracted, thereby restoring a partial vacuum in the chest cavity and inflating the lungs. Protoype devices were tested extensively in a benchtop apparatus to confirm operation within the physiologically relevant parameters and also to ensure operation in a fluid environment. After these tests were concluded, a series of implants into a rabbit was performed under a nonsurvival protocol. These initial implants served to further develop the surgical implantation procedure and provide feedback to the engineering design team. After additional design iteration and testing, a device was implanted into a dog, again under a nonsurvival protocol. The implant prior to closure of the surgical site is shown in Figure 21.6. After wound closure, the partial vacuum in the chest cavity was restored and ultrasound video was obtained of the device functioning (Figure 21.7). The device was operated intermittently for approximately two hours, showing the expected reversible deformations similar to the natural muscle, at which time the device failed and the approved protocol required termination of the experiment. These experiments point the way towards the eventual use of EPAM devices as implantable functional replacements for natural muscle. As reported, further developments will be directed at improving packaging for long term operation in a fluidfilled implantable device environment and, for the diaphragm application specifically, increasing displacement volumes to achieve acceptable levels of gas exchange in the lungs.

404

Biomedical Applications of Electroactive Polymer Actuators External Internal

Phrenic Nerve

Diaphragm

External Inductive Power Supply

Induction Coil (a)

EPAM Active Area

Suture Fabric Rigid Silicone Frame

Electrical Connection (b)

Figure 21.5 Schematic overview (a) of the EPAM artificial diapragm concept and (b) initial prototype design (Reprinted with permission from ASM Proceedings of MPMD 2007, Medical Device Applications of Dielectric Elastomer Based Artificial Muscles by Bashkin, J.S., Heim, J., Prahlad, H. et al. Copyright (2007) ASM). Electrical Lead

EPAM Active Area

Rigid Frame

Figure 21.6 Surgical implantation of the artificial diaphragm (Reprinted with permission from ASM Proceedings of MPMD 2007, Medical Device Applications of Dielectric Elastomer-Based Artificial Muscles by Bashkin, J. S., Heim, J., Prahlad, H. et al. Copyright (2007) ASM).

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Figure 21.7 Ultrasound image of the fully implanted artificial diaphragm device, viewed edge on. The circle shows the region of the artificial diaphragm; the arrow shows the remainder of the natural diaphragm muscle. (Reprinted with permission from ASM Proceedings of MPMD 2007, Medical Device Applications of Dielectric Elastomer-Based Artificial Muscles by Bashkin, J. S., Heim, J., Prahlad, H. et al. Copyright (2007) ASM).

21.5 Implantable Artificial Facial Muscles Another research area with a great deal of potential is the use of EPAM actuators for facial reconstruction. One application currently in development [5] is the implantation of EPAM for eyelid reanimation. Damaged and paralyzed eyelids caused by burns or trauma (perhaps sustained during combat by soldiers) or illness (such as cranial nerve inflammation, stroke or head and neck tumors) present a significant risk to the affected person’s eyesight. An eyelid that cannot close fully (lagophthalmos) results in continuous exposure of the eye meaning less protection from foreign bodies, abrasions and drying. Corneal exposure eventually leads to eye desiccation, ulceration, perforation and blindness. Lagophthalmos is especially a problem for eyelids after thermal or chemical burns. Second and third degree burns cause contraction of the eyelid and the shortened eyelid often cannot close completely. In addition to the performance characteristics discussed previously, EPAM has additional properties that make it especially well suited for this reconstruction application. Since there is no continuous electrical current flow through the polymer, EPAM is efficient at the relatively low frequency operation required in this and many other implantable applications. Thus, minimal power is consumed, translating directly to a longer battery life and less recharging of the implanted device. Since virtually no heat is generated by the EPAM actuator, there is negligible risk of thermal damage to surrounding tissue. Additionally, EPAM performs silently. This is especially important for the present application where an actuator would be implanted in the temple in close proximity to the ears. Finally, the mechanically compliant nature of EPAM films combined with their high power density gives them significant advantages for this application over motors, electromagnetic coils and piezoelectric ceramics. A possible implementation of an eyelid actuation system is shown in Figure 21.8. An EPAM module, implanted in the temple would drive a sling mechanism to open and close an eyelid.

406

Biomedical Applications of Electroactive Polymer Actuators

Figure 21.8 Model of human skull with a possible implementation of an implantable EPAM actuator for eyelid reanimation. The arrow indicates the EPAM module with slings for the upper and lower eyelids (Reprinted with permission from Laryngoscope, Restoration of Eyelid Closure in Facial Paralysis Using Artificial Muscle: Preliminary Cadaveric Analysis by Tollefson, T.T. and Senders, C.W, 117, 11, 1907-11. Copyright (2007) Lippincott, Williams and Wilkins).

21.6 Limb Prosthetics and Orthotics Electroactive polymer actuators, and EPAM actuators in particular, are often called ‘artificial muscles’ because of the similarity in strain and force to natural skeletal muscle. Perhaps no other actuator application conveys the notion of an ‘artificial muscle’ more than those for moving natural or prosthetic limbs. Application of EPAM for limb prosthetics and orthotics still requires significant further development of the EPAM technology in order to address issues in scaling up stroke and force, as well as packaging issues related to reliability and safety at relatively high power levels (compared to most other applications discussed here). However, the fundamental demonstrated performance of EPAM in smaller devices shows the promise of this ‘artificial muscle’ technology. The performance on a per mass and per volume basis demonstrated in smaller devices is sufficient to create actuators that can replace or augment the function of natural muscle. Other characteristics of EPAM that make EPAM well suited for such limb prosthetic and orthotic applications include:

 Quiet operation – noise from prosthetic limbs and orthotics can be annoying and draw unwanted attention to the wearer.

 High compliance – EPAM is soft and compliant like natural muscle. Such compliance means that EPAM can better duplicate the behavior of natural muscle-driven limbs when interacting with objects (e.g. holding delicate objects or responding to accidental impacts or unknown disturbances). Just as the stiffness and damping of natural muscle

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can be modulated using neural and other feedback loops [6], the mechanical impedance of EPAM can also be modulated with simple electrical driving circuits [7].  Similar size and shape to natural muscle – EPAM can be used in a direct drive mode where prosthetic limbs incorporate actuators that are of similar size and shape to natural skeletal muscle, allowing the entire limb to appear and function like a natural limb. Some proof-of-principle devices have been demonstrated and specific promising applications identified. An EPAM artificial muscle actuator installed on a full-size plastic model of a human skeleton arm is shown in Figure 21.9. The actuator is acting much like a bicep. The artificial bicep is a rolled actuator, made by scrolling EPAM material into a cylindrical shape. Upon application of a voltage, the cylinder elongates. This particular muscle was capable of about 30% strain and about 2 kg of force. This mostly hollow muscle was not packaged for practical use nor was the force and lifetime sufficient for a practical elbow actuator. Nonetheless, this actuator conveys the promise of EPAM for prosthetics.

Figure 21.9 Still frame from a video of an EPAM ‘bicep’ actuator (rolled actuator) moving the elbow joint of a human skeletal model (Used by permission of the authors. All rights reserved).

A somewhat more refined demonstration arm was built by ADA Technologies (Littleton, Colorado, USA) using prototype EPAM actuators made by SRI International. This arm, shown in Figure 21.10, was intended as a test-bed for advanced prosthetic concepts. The arm again uses rolled actuators in place of natural muscle but it uses a ratchet mechanism to augment the stroke and force of the individual actuators. The actuators go through several cycles of elongation and contraction in order to produce the desired motion. This approach is not as desirable as making a larger direct-drive actuator

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Figure 21.10 Prosthetic arm test-bed powered by EPAM; (top) a bundle of artificial muscles installed in a prototype of a prosthetic arm (each muscle is about 10 cm long); (bottom) the artificial muscle powered arm shown lifting a weight (Reproduced by kind permission of ADA Technologies, Inc.).

that can exploit the compliance and smooth quiet operation of EPAM but it does allow for short term exploration of the use of artificial muscle in limb prosthetics. In 2007, MIT professor Hugh Herr demonstrated the first powered ankle prosthetic device [8]. Mechanical compliance is one key to the proper operation of an ankle prosthetic since the foot is constantly interacting with surfaces of varying compliance. The MIT ankle does not use EPAM but rather uses a linear servomotor drive (rotary motor plus a lead screw) in combination with a spring in order to achieve a controllable compliance. A similar ankle orthotic device has also been developed [9]. Replacing this relatively heavy, bulky and noisy mechanism with an EPAM actuator would be desirable. Proper management and recovery of the elastic energy of the musculature around the ankle joint is important for an efficient and natural appearing walking gait. In the case of a natural ankle, approximately three-quarters of the energy expended during the gait cycle is recovered elastically [8]. While the maximum force requirements of lower limb muscles are high, normal walking on level ground does not require large energy expenditure by the muscles. Thus, even a relatively small EPAM actuator that effectively uses its elastic compliance in addition to actuation can offer significant advantages for a lower limb prosthetic or orthotic in typical usage [10].

21.7 Mechanical Actuation for ‘Active’ Cell Culture Assays Studies of tissues grown in cell culture are fundamental to experimental cell biology. Typically, these studies are performed by growing adherent cells in plastic wells that provide a rigid surface for attachment and cellular organization. The in vivo environments for cells that make up many tissue types is mechanically dynamic, however, and it is difficult to study the impact of mechanical stress on the behaviour of cells. Cell types of interest include stem

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cells, cardiac cells, cells involved in scar tissue formation and smooth muscle cells. SRI, in partnership with collaborators, has been developing a new platform for studying adherent cells that would allow for the systematic and uniform imposition of static or cyclical mechanical stress on cells. The hypothesis behind this work is that by growing cells on a compliant substrate that can be actuated, data on cell behaviour and drug toxicity, for example, might be obtained that is more physiologically relevant than data available to date. Previous studies involving the phasic stretching of myocytes by methods that are more conventional have shown a beneficial effect on the growth of the heart cells [11]. To achieve this goal, we developed dielectric elastomer based systems that are compatible with cell adhesion and growth and operate functionally in the presence of growth media and the warm, humid environment found in a cell incubator. Cells are stretched by approximately 15 % using a planar dielectric elastomer actuator (Figure 21.11). In the future, we anticipate an integrated system in which cells can grow actuate directly on top of a micropatterned dielectric elastomer device. This would permit a simpler and more controllable cell culture assay than is possible using conventional electromagnetic actuation. The emphasis is on creating a system in which the mechanical strain is controllable and uniform across the entire experimental cell population, and which enables a relatively high experimental throughput. Actuation using the EPAM has been shown to be a viable approach to inducing phasic strains in the cells.

Electroded area coated with cells for culture

Patterned Electrode Voltage Off

Voltage On

Elastomeric EPAM membrane

Figure 21.11 Schematic illustrating the implementation of active heart cell stretching using planar EPAM actuators. Cells are grown on a flexible membrane actuated with phasic stretching using EPAM. Depending on EPAM and electrode structure, cells can be stretched in either planar direction or in both directions at the same time.

21.8 Conclusions As EPAM technology has evolved into commercially available products, the performance characteristics have also improved dramatically, with optimizations of the dielectric elastomer materials, manufacturing quality and automated production processes. The applications for the technology will also continue to expand from drop-in replacements

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for traditional valve and pump components into more exotic uses such as implantable replacements for lost muscle function and as a platform for biology research. Many of these applications share common features, such as an interaction between EPAM films and biological tissue or exposure to moist environments. For other applications such as drug delivery pumps, basic performance characteristics of force, stroke and lifetime must continue to improve incrementally before EPAM based products are commercially viable. The unique combination of performance attributes exhibited by EPAM, however, lends the technology to a host of biomedical applications that are largely limited only by imagination.

References 1. Pelrine, R., Kornbluh, R., Pei, Q. and Joseph, J. (2000) High-speed electrically actuated elastomers with over 100% strain, Science, 287: 836–9. 2. Prahlad, H., Pelrine, R., Kornbluh, R., et al. Programmable surface deformation: thickness-mode electroactive polymer actuators and their applications, in Smart Structures and Materials 2005: Electroactive Polymer Actuators and Devices (EAPAD), Proceedings of SPIE (ed. Bar-Cohen, Y.), 5759, San Diego, CA., 6 May 2005, 102–13. 3. Priplata, A., Niemi, J. B., Harry, J. D., et al. Vibrating insoles and balance control in elderly people, Lancet 2003; 362: 1123–4. 4. Bashkin, J. S., Heim, J., Prahlad, H., et al. (2007) Medical Device Applications of Dielectric Elastomer-Based Artificial Muscles, Medical Device Materials IV: Proceedings of the 2007 Materials and Processes for Medical Devices Conference, Palm Desert, CA. 5. Tollefson, T. T. and Senders, C. W. (2007) Restoration of Eyelid Closure in Facial Paralysis Using Artificial Muscle: Preliminary Cadaveric Analysis, Laryngoscope, 117, 1907–11. 6. Winters, J., Muscle as an Actuator for Intelligent Robots, Robotics Research: Trans. Robotics International of SME, Scottsdale, AZ, 18–21 August 1986, MS-86-76-1–18. 7. Pelrine, R. (2008) Variable stiffness mode: devices and applications, in Dielectric Elastomers as Electromechnical transducers (eds Carpi, F., DeRossi, D., Kornbluh, R.), Elsevier, Oxford, UK, 141–5. 8. Au, S. K. and Herr, H., Initial experimental study on dynamic interaction between an amputee and a powered ankle-foot prosthesis, Workshop on Dynamic Walking: Mechanics and Control of Human and Robot Locomotion, Ann Arbor, MI, May 2006. 9. Blaya, J. and Herr, H., Adaptive Control of a Variable-Impedance Ankle-Foot Orthosis to Assist Drop-Foot Gait, IEEE Trans. Neural Systems & Rehabilitation Engineering, 12(1), 24–31, 2004. 10. Herr, H. and Kornbluh, R. (2004) New Horizons For Orthotic And Prosthetic Technology: Artificial Muscle For Ambulation, Smart Structures and Materials: Electroactive Polymer Actuators and Devices (ed. Bar-Cohen, Y.), Proceedings of SPIE, 5385, 1–9. 11. Fink, C., Su¨leman, E., Kralisch, D., et al. Chronic stretch of engineered heart tissue induces hypertrophy and functional improvement, FASEB Journal, 14, April 2000, 669–79.

22 MRI Compatible Device for Robotic Assisted Interventions to Prostate Cancer Jean-Se´bastien Plante1, Lauren Devita2, Kenjiro Tadakuma2 and Steven Dubowsky2 2

1 Universite´ de Sherbrooke, Canada Massachusetts Institute of Technology, USA

22.1 Introduction The ability to perform cancer therapies within the bore of Magnetic Resonance Imaging (MRI) systems has the potential to greatly improve cancer survival rates and quality of life after treatment [1]. For example, a real-time image and a MRI compatible manipulator would allow surgeons to accurately guide a needle to a malignant tumor while avoiding vital structures. Very small tumors (<3 mm) can be localized and reached enabling early cancer detection and local treatment. Health risks, treatment costs and side effects are minimized. Applications of MRI guided robotic surgery extend beyond cancer therapy and could have significant impact in neurology, orthopedics and cardiology. There has been substantial research in the development of in-bore MRI manipulation for diagnostic and surgical procedures, as discussed below. These manipulator systems must be compatible with the very high magnetic fields (~3 Teslas) used in MRI. Nonmagnetic materials and low electrical currents must be use to assure safety and prevent interference with the imaging process. Conventional robotic actuators, which are typically electromagnetic devices, must be more than 1 m away from the center of the magnet [2]. Hence, they require complex, expensive, and cumbersome transmissions making this approach unpractical. Dielectric Elastomer Actuators (DEAs) are a promising alternative for MRI manipulation because they can be actuated directly within the bore of the MRI without affecting image quality [3]. Unlike other MRI compatible actuators, such as piezoelectrics and pneumatics, DEAs are an attractive solution because they offer good mechanical performance while

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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being simple and inexpensive (Figure 22.1). However, using DEAs in practical robotics and mechatronics systems is not easy due to their time-dependent failure modes that limit their performance and reliability. Recent studies have shown that using DEAs intermittently at high speeds can significantly increase their performance and reliability [4].

Figure 22.1 Dielectric elastomer actuator under 100 % extension.

Binary actuation is a design paradigm that exploits the advantages of using DEAs intermittently. Binary systems are the mechanical equivalent of digital electronics [5]. A binary system as intended here is driven by bistable actuators that switch between two possible states, extended or retracted. Overall system architecture is greatly simplified compared to conventional robotic systems since very few sensors are needed [6]. Further, in the context of MRI applications, bistable DEAs can advantageously be turned off before an image is taken, further reducing the chances of affecting the image. This chapter presents a binary manipulator for MRI guided prostate cancer biopsies and brachytherapies using a transperineal approach (Figure 22.2). The problem of prostate cancer detection and treatment is exposed along with a review of alternative approaches. The proposed manipulator concept is presented and an analytical model of its performance is developed. Model predictions are compared with experimental data. Results show that Bistable Bistable Assembly Assembly

Needle

Needle Needle

(a) System Concept

Figure 22.2

(b) Experimental Setup

Elastically averaged parallel manipulator for MRI guided prostate cancer therapy.

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the proposed binary manipulator concept has the potential to meet the clinical requirements of transperineal needle insertion. The proposed concept is effective and could easily extend to general manipulation problems.

22.2 Prostate Cancer Therapy Prostate cancer is the most frequently diagnosed cancer in men and the number two cause of cancer related death in men after lung cancer [7]. In the United States alone, 218 890 men were diagnosed with prostate cancer in 2007 while 30 000 men die from the disease each year [7, 8]. The number of deaths has been declining since the early 1990s due to better detection and treatment methods. 22.2.1

Prostate Cancer Detection

To date, needle biopsy is the best method to confirm if a malignant prostate tumor is present. However, studies show that current biopsy techniques miss up to 20 % of all prostate cancers, leaving many men with ‘false negative’ results [9, 10]. In these cases, the tumor will go undetected. Left untreated, it may be too late to successfully treat, especially if the cancer has spread beyond the prostate. Accurate biopsies are vital to prostate cancer detection and, hence, treatment. Early stage detection requires accurate imaging systems and precise needle insertion techniques to probe small potential tumors. Currently, ultrasound imaging is used to manually insert a biopsy needle to the prostate. Unfortunately, the biopsy needle often misses the tumor because ultrasound images are too low resolution to see small, early stage, tumors. Tumors smaller than 5 mm are not detected by ultrasound imaging [11]. Only about 20 % of tumors between 5 and 10 mm are detected by ultrasound. Even large tumors, of the order of 20–25 mm, are only detected 79 % of the time. Ultrasound images, therefore, cannot be used to detect or treat millimeter size prostate tumors. Figure 22.3 shows ultrasound and MR images of the same prostate [12]. Even to the untrained eye, it is clear that MR images are far more detailed than ultrasound images.

Figure 22.3 Ultrasound image (left) and MR image (right) of same prostate (Reproduced with permission from Real time MRI-ultrasound image guided stereotactic prostate biopsy by Kaplan, I., Oldenburg, N.E., Meskell, P., et al. Magnetic Resonance Imaging, 20, 3. Copyright (2002) Elsevier).

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Given the unique imaging capabilities of MRI, it is possible to diagnose and treat small, millimeter size tumors that cannot be detected otherwise [13, 14]. 22.2.2

Prostate Cancer Treatment

In the past, radical prostatectomy has been the only option to treat prostate cancer. Important undesirable side effects such as incontinence and impotence combined with long treatment duration and recovery time has caused many men to turn to new treatments such as radiation therapy. Brachytherapy is an interesting form of radiation therapy because it is an outpatient procedure needing a recovery time of only a few days. During brachytherapy, similar to needle biopsy, a needle is inserted into the tumor in the prostate, typically using an ultrasound image for guidance (Figure 22.4). The needle deposits small radioactive pellets in the tumorous region. Cryotherapy, where very low temperature fluids such as liquid nitrogen are injected to freeze a tumor, can also be used. Figure 22.4 shows a transperineal approach which, compared to the other alternative of transrectal approach, has lower risks of hitting vital structures and does not require antibiotics because the insertion avoids penetrating the bowels. Needle, delivering seeds into prostate

Catheter in urethra

Ultrasound probe in rectum for needle guidance

Template to aid accurate placement of the needles delivering the seeds

Figure 22.4 Prostate cancer transperineal brachytherapy treatment (The Prostate Cancer Centre, www.prostatecancercentre.co.uk).

Transperineal brachytherapy is growing in popularity over other treatment options because of a significantly smaller chance of incontinence and impotence, quick treatment and recovery and high success rates [15]. As brachytherapy becomes the method of choice, the procedure must become more reliable. Early detection through accurate biopsy needle placement is crucial to treat prostate cancer. To increase prostate cancer survival rates, it follows that needle placement methods must be improved with MRI guided methods.

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22.2.3

415

Needle Placement in MRI Systems

Robotic needle manipulators have been developed for assisted interventions in open-bore MRI [16, 17], which uses split magnets that leave room for large remote manipulators to be used. Compared to closed-bore MRI, open-bore MRI operate at lower magnetic fields (1.5 versus 3.0 Teslas) and has significantly lower resolution. Image quality of open-bore MRI is not detailed enough or suitable for accurately treating a tumor on the order of 5 mm [18]. Closed-bore MRI manipulation requires actuators with excellent MRI compatibility because of their proximity to the image center. Piezoelectric motors have been reported to cause image distortion in closed-bore MRI [19]. Piezoelectric motors are also relatively slow, complex and expensive. Manipulators using pneumatics have had limited success in conjunction with closed-bore MRI [20, 21]. The main drawbacks to this approach are control issues due to the imprecision and compliance of pneumatics. Recent advances have been made using pneumatic stepper motors to eliminate these effects [22]. However, pneumatic systems, using continuous cylinders or step motors, tend to be complex and expensive. Compared to other actuation alternatives, DEAs have a huge potential for closed-bore MRI manipulation. They have excellent MRI compatibility, offer good performance and are simple and low cost. The binary manipulator system presented in the next section is an illustration of a potential future practical MRI manipulator using DEAs.

22.3 Elastically Averaged Parallel Manipulator Using Dielectric Elastomer Actuators The manipulator is intended to be used inside the bore of a closed-bore MRI machine to perform transperineal prostate cancer needle biopsy, brachytherapy and cryotherapy (all can be performed by inserting a needle through the perineum). This section discusses design requirements, presents the manipulator concept and proposes an analytical model of the manipulator. 22.3.1

Design Requirements

The following design requirements for system size, workspace, precision and forces have been developed in collaboration with doctors and researchers at Brigham and Women’s Hospital in Boston, Massachusetts, USA. 22.3.1.1

Size and Workspace

The device must be MRI compatible and able to fit between the patient’s legs while inside the bore of the MRI machine (Figure 22.5).The bore of a typical MRI machine is 550 mm in diameter. The patient’s legs are propped up to provide access to the perineum. This leaves a small space for the device, requiring that the device be no larger than a 200 mm diameter cylinder, 500 mm deep. Under the control of a doctor using a real-time MR image, the proposed device is required to reach a target (a tumor) in the prostate by penetrating the perineum as shown in Figure 22.6. The average prostate is located 60–80 mm from the perineum, and the size of the average prostate is 30–50 mm in the z-direction [23]. The needle must be able to travel between 60 and 130 mm in the z-direction. The required workspace is an 80  70 mm elliptical cylinder protruding by 70 mm in the z-direction, as shown in Figures 22.5 and 22.6. This workspace is

416

Biomedical Applications of Electroactive Polymer Actuators Bore

Device

Workspace Prostate Workspace

Device

80 mm

Prostate

Cradle 70 mm

70 mm 500 mm

200 mm

Figure 22.5 Size and workspace requirements.

Workspace Perineum

Prostate

70 mm

y 80 mm

x z Needle Guide

Rectum 60 mm 70 mm

Figure 22.6

Needle path.

larger than the average normal prostate, but is necessary to accommodate the frequently enlarged prostates of cancer patients as well as position differences from patient to patient. 22.3.1.2

Precision

Binary systems have discrete workspaces consisting in all possible end-effector locations. Here, precision is defined as the minimum distance between a random point in the required workspace from the closest possible needle tip location. Manual needle insertion methods currently in use have a perforated template with holes 5 mm apart to guide the needle.

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Statistical studies showed that the average distance of these templates is 1.9 mm [24]. The target for the manipulator is to match this level of precision or do better. It worth mentioning that the manipulator will be guided by a surgeon using real-time MR images. In this case, the absolute precision becomes less critical since corrections can be applied during insertion. 22.3.1.3

Forces

Tests were performed on beef muscle to establish the magnitude of the external forces applied on the needle during insertion [24]. Beef muscle has similar properties to those of the tissue between the perineum and the prostate. In these tests, forces are considered to be applied at the needle’s point of entry. Combining these results with those of the literature give a maximum axial force of 14 N and a maximum radial force of 1.6 N for a beveled needle tip [25]. A trihedral needle tip would lower the axial force to ~8.25 N and the radial force to ~0.5 N. Hence, trihedral needles are desirable. 22.3.2

Manipulator Concept

A schematic representation of the proposed manipulator is shown in Figure 22.7. It is an elastically averaged parallel manipulator having two planes, each containing a circular array of six bistable actuators and six spring elements. A biopsy needle runs through a rigid tube attached at the center point of each plane and advances from Plane 1, through Plane 2 to the target. The planes are separated by a distance p and the needle tip is at a distance zd from the first plane. Needle manipulation is achieved by changing the location of the center point of each plane when the static equilibrium of the springs is modulated by the bistable modules. Compared to conventional serial chain manipulators, the proposed parallel architecture provides higher system stiffness and higher precision.

k6 k1

k5

Center Point

Plane 1 Needle Tip

Plane 2

k2

k4 k3

Needle

p

Bistable Module

Node zd

(a) One plane of the device

Figure 22.7

(b) Two plane assembly

Schematic of the proposed manipulator.

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The bistable module concept is shown in Figure 22.8. Bistability is achieved by flipping a bistable truss with two independent cone-shaped DEAs. The module output has a stroke z between its ‘extended’ and ‘retracted’ states (Figure 22.7 shows all bistable assemblies in the ‘retracted’ position). Using DEAs in a bistable fashion has a considerable advantage over continuous operation because it increases actuator performance and life [4]. Moreover, power is not needed to hold a state which reduces energy consumption and the risk of interferences with the MRI process.

Cone DEA 1

Bi-stable truss

Module Output

ζ

Cone DEA 2

(a) Concept

Figure 22.8

22.3.3

(b) Prototype

Bistable module.

Manipulator Analytical Model

A binary manipulator input, Q, can be represented in a binary sequence of 0 and 1: Q ¼ ½ a1

a2

...

an Š

ð22:1Þ

where ai ¼ 1 or 0 is the state (extended or retracted) of the ith bistable module and n is the total number of bistable modules. Binary systems can only reach a finite set of discrete points (2n points) covering the system’s workspace. In most systems, these discrete points are not evenly distributed in the workspace. Hence, an analytical model that maps the manipulator workspace for any given inputs is essential to optimize the manipulator design for a given task. For a given set of inputs, the spatial configuration of the elastically averaged parallel manipulator is found by resolving the static equilibrium in each plane of the device. Typical spring deformations and applied forces in a single plane of the device are represented schematically in Figure 22.9; only three springs are shown for clarity. The model uses the following input parameters: ki 10i ¼ l0iwi bi ¼ iwi di ¼ iwi

fext W

stiffness of ith spring position vector of the ith undeformed spring pre-stretch vector of the ith spring. stroke vector of ith bistable module. The magnitude of the stroke vector is given by i ¼ ai with ai (0 or 1) the state of the bistable module and  the stroke. external force vector at center point weight vector at center point

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y x F1

k2

F2

k1 u

Undeformed Center Point

Deformed Center Point

ki

l0i

W fext

li

βi

Fi

δi (a)

(b)

Figure 22.9 Analytical model variable definition: (a) node and center point displacements; (b) free body diagram.

to compute the following output parameters: u ¼ ue li ¼ ltvi F i ¼ Fi vi

displacement vector of the center point position vector of the ith deformed spring internal force vector of the ith spring

In the preceding lists, wi,vi and e are the respective unit vectors of the undeformed direction of the ith spring, the deformed direction of the ith spring and the displacement of the center point. For static equilibrium, the sum of the forces at the center point must equal zero: X X forces ¼ Fi þ f ext þ W ¼ 0 ð22:2Þ i

The internal force of the ith spring generated by the spring elongation is given by: Fi ¼ ki ðli

l0i Þvi ¼ ki ½ðl0i þ i þ i Þwi

ue

l0i vi Š

ð22:3Þ

Equations (22.2) and (22.3) are solved iteratively for the displacement vector of the center point, u. The method is used to find the center point locations of each plane, u1 ¼ (x1,y1,z1) and u2 ¼ (x2,y2,z2), from which the needle tip (end-effector) location can be found by: ud ¼ ðxd ; yd ; zd Þ ¼ u1 þ

zd ðu2 p

u1 Þ

ð22:4Þ

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where p and zd are geometric parameters defined in Figure 22.7. In these calculations, the origin is taken as the center point of Plane 1 before any perturbation. The model is used to compute the full manipulator workspace corresponding to all combinations of the bistable modules states. The manipulator system stiffness represents its ability to resist external forces without major deflections. System stiffness is defined by the equivalent stiffness at the tip of the manipulator: Keq ; tip ¼

keq1 keq2 p2 keq1 ðzd Þ2 þ keq2 ðzd

ð22:5Þ

pÞ 2

where keq1 and keq2 are the equivalent stiffness in Planes 1 and 2, p and zd are geometrical parameters defined in Figure 22.7 [24]. Equation (22.5) is defined in a radial plane of the device and must be evaluated from 0 to 360° to evaluate the manipulator stiffness in all directions. If the effects of external forces, fext, and gravity, W, can be neglected, the displacement vector of the center point of each plane can be obtained from Equations (22.2) and (22.3) as: 1 u¼ X

ki

X

ki ½ðl0i þ i þ i Þwi

l0i vi Š

ð22:6Þ

i

i

Defining a dimensionless ratio of the spring stiffness, i¼ ki/ k1, yields: 1 u¼ X

i

X

i ½ðl0i þ i þ i Þwi

l0i vi Š

ð22:7Þ

i

i

Equation (22.7) shows that when there are no external forces or gravity, the displacement vectors of the center points are functions the ratio of the spring stiffness, not their absolute value. Hence, the manipulator’s workspace is independent of its overall stiffness and a very ‘soft’ manipulator can have the same workspace as a very ‘stiff’ one.

22.4 Results The laboratory prototype of Figure 22.2b was developed to evaluate the effectiveness of the manipulator concept. The key design parameters are listed in Table 22.1. The system has an outer diameter of 400 mm and a length of 450 mm. It uses 12 functional bistable modules, each using a pair of cone-shaped DEAs. These cone actuators are hand fabricated using two active polymer layers made from 1.5 mm thick films of 3M’s VHB4905/4910. The cone actuators have an outer diameter of 100 mm, develop a maximum force of 6 N and have a stroke of 12 mm. The prototype was designed using the performance specifications of current handfabricated DEAs. Hand fabrication limits the number of active film layers in each actuator

MRI Compatible Device for Robotic Assisted Interventions to Prostate Cancer Table 22.1

421

Manipulator parameters Dimensions (mm)

p

l0i

i

zd

180

86

13

560–630

Spring Constants, k ¼ 0.044 N/mm

Plane 1 Plane 2

k1

k2

k3

k4

k5

k6

3k/2 k/2

k/2 k

k k/2

k/2 k/2

k k/2

k/2 k

resulting in relatively large actuators with limited forces. Consequently, the prototype had to be designed at twice the scale than the clinical device and the spring’s stiffness had to be lowered to prevent actuator saturation. Developing smaller and stronger DEAs using many active layers is feasible through the development of appropriate actuator manufacturing techniques (currently under development). 22.4.1

Analytical Results

The analytical model developed in Section 22.3.3 is used to study the prototype workspace, precision and stiffness properties. Figure 22.10a shows the needle workspace at penetrations from 60 to 130 mm. Figure 22.10b shows the cross-section of the workspace at a penetration of 110 mm (zd ¼ 610 mm). As discussed above, a binary system’s precision is largely determined by the number of actuators. This system with 12 bistable modules can

(a)

(b)

Figure 22.10 Analytically predicted workspace: (a) full needle workspace; (b) workspace at an insertion depth of 110 mm.

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reach 212 ¼ 4096 points. In general, these points are not distinct and uniformly distributed over the workspace (Figure 22.10b). Here, the analytical model was used to select the spring constants in order to eliminate system symmetries and associated redundant points. As a result, the workspace of Figure 22.10 is not symmetric. The workspace is relatively well distributed with an increasing point density toward the center of the workspace, within the prostate, where tumors are most likely to be found. Manipulator precision is evaluated from 1000 random target points within the workspace. The minimum distance between these points and the closest possible needle tip (end-effector) position was calculated. The manipulator must be able to reach within a distance of 1.9 mm of any point in the prostate. The error distribution is shown in Figure 22.11a. The average distance and standard deviation for a random point to the nearest possible end-effector point in the required workspace are 0.68 mm and 0.51 mm, respectively. The same values for a prostate sized workspace are 0.41 mm and 0.21 mm. Clearly, the manipulator design meets its precision requirements with only 12 binary inputs. This is not surprising since the resolution of binary systems increases drastically with the number of inputs. For example, doubling the number of binary inputs to 24, which is technically and economically feasible with DEAs, would drop the minimum distance in the micrometer range.

350 300 Number of Trials

90

Required Workspace Average Prostate

250

120

0.015

60

0.01

150

200

30

0.005 180

150

0

100

0 –1

330

210

50 0

1

2 3 Distance (mm) (a)

4

5

6

300

240 270 (b)

Figure 22.11 Analytically predictions: (a) minimum distance from needle tip to target at an insertion depth of 110 mm; (b) system stiffness at needle entry point (N/mm).

Figure 22.11b shows the distribution of system stiffness in the radial direction (perpendicular to the needle) at the needle entry point (zd ¼ 500 mm). The figure shows that system stiffness is quite uniform although absolute values are low. Recall that, for this prototype, system stiffness is limited by the low forces of the handmade DEAs. Developing higher stiffness requires actuators with higher force output. To meet the clinical requirements on needle insertion forces, it is estimated that the number of active layers must increase from 2 to 30, which is technically feasible with proper actuator manufacturing techniques [24].

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22.4.2

423

Experimental Results

The motion of the needle tip was measured experimentally and compared with predictions of the analytical model. The tip motion is measured with a magnetic position sensor (miniBIRD model 80 from Ascension Technology). This electromagnetic tracking system measures the three positions and three orientations of a small sensor (1.3 mm in diameter, 6.5 mm in length) with RMS accuracies of 1.4 mm and 0.5°. The sensor is small compared to the manipulator and does not affect its performance. The measured location of 54 random inputs over the system’s 4096 possible inputs is shown in Figure 22.12a. The general size and shape of the workspace is essentially the same as that predicted by the analysis (Figure 22.10b). Analytical predictions of the needle tip location with experimental measurements are compared in Figure 22.12b. The circle shows the size of an average prostate. In this area, the average distance between the simulated and experimental points is about 3 mm, which is smaller than the required precision of þ/–5 mm. It should be noted that the sensor itself has an RMS error of about 1.4 mm, suggesting that more precise laser measurements and calibration would significantly reduce the experimental errors. Also, the prototype is a laboratory device containing many geometric error sources that would easily be eliminated in a production environment.

40

Required workspace Average prostate Experimental Theoretical 15

30 20

10 10

–40

–30

–20

5

0 –10 0 –10

10

20

30

40 –25

–15

0 –5 –5

5

15

25

–10 –20

–15

–30 –40

(a)

(b)

Figure 22.12 Experimental measurements of needle tip position (mm): (a) 54 random points; (b) comparison with analytical model.

Finally, the MRI compatibility of the prototype was verified in a 3-Tesla MRI at Harvard’s Brigham and Woman’s Hospital [24]. These tests confirmed previous results on individual actuators showing that DEA technology is MRI compatible [3]. Specifically, the manipulator was not degraded by the high magnetic fields and the MR images were not affected by the manipulator using conventional filtering.

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22.5 Conclusions This chapter presented an MRI compatible manipulator concept based on elastically averaged binary DEAs for prostrate cancer detection and treatment. A manipulator design using 12 binary degrees-of-freedom was proposed along with an analytical model of its performance. Experimental and analytical validations suggest the approach to have sufficient precision and workspace range to meet the medical requirements. However, system size and stiffness are currently limited by the low forces of the hand-fabricated DEAs available in this study. Appropriate manufacturing techniques for high force DEAs are currently under development. Tests performed in a 3-Tesla MRI machine confirmed the excellent MRI compatibility of the technology. Results presented in this chapter suggest that the proposed manipulation approach based on elastically averaged parallel manipulation is an effective way of using DEAs in practical robotics and mechatronics systems. Other medical procedures requiring precise robotic needle placement in MRI environments could beneficiate from this work, such as breast cancer detection and treatment, endovascular surgeries and spinal procedures. Possible applications of elastically averaged parallel manipulation using bistable DEAs extend to general robotic tasks, inside or outside MRI systems, providing precision, simplicity and low costs.

Acknowledgements The important contributions of Dr Daniel Kacher, Dr Joseph Roebuck and Dr Simon DiMaio of Harvard’s Brigham and Women’s Hospital (BWH) are truly appreciated.

References 1. Jolesz, F. (2005) Future perspectives for intraoperative MRI, Neurosurgery Clinics of North America, 16, 201–13. 2. Chinzei, K., Kikinis, R. and Jolesz, F. (1999) MR Compatibility of Mechatronic Devices: Design Criteria, Medical Image Computing and Computer-Assisted Intervention, 1020–30. 3. Vogan, J., Wingert, A., Hafez, M., et al. (2004) Manipulation in MRI Devices Using Electrostrictive Polymer Actuators: with an Application to Reconfigurable Imaging Coils, Proceedings of the 2004 IEEE International Conference on Robotics and Automation, New Orleans, LA, 23 April–2 May, 2498–504. 4. Plante, J.S. and Dubowsky, S. (2006) Large-Scale Failure Modes of Dielectric Elastomer Actuators, Int. J. Solids and Structures, 43, 7727–51. 5. Roth, B., Rastegar, J. and Scheinman, V. (1973) On the Design of Computer Controlled Manipulators, Proceedings of the First CISM-IFToMM Symposium on Theory and Practice of Robots and Manipulators, 93–113. 6. Sujan, V., Lichter, M. and Dubowsky, S. (2001) Lightweight Hyper-redundant Binary Elements for Planetary Exploration Robots, Proceedings of the 2001 IEEE/ASME International Conference on Advanced Intelligent Mechatronics (AIM ‘01), Como, Italy, 8–12 July 2001. 7. Jemal, A., Murray, T., Ward, E., et al. (2005) Cancer Statistics 2005, C.A.: A Cancer Journal for Clinicians, 55, 10–30. 8. Jemal, A. et al. (2007) Cancer Facts and Figures 2007, American Cancer Society.

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9. Babaian, R.J. et al., NCCN Clinical Practice Guidelines in Oncology: Prostate Cancer Early Detection, VI, available online from the National Comprehensive Ccancer Network www.nccn.org, 2006. 10. Stoumbakis, N., Cookson, M.S., Reuter, V.E. and Fair, W.R. (1997) Clinical significance of repeat sextant biopsies in prostate cancer patients, Urology, 49 (3A Suppl), 113–8. 11. Jamis-Dow, C.A., Choyke, P.L., Jennings, S.B., et al. (1996) Small (S3-cm) Renal Masses: Detection with CT versus US and Pathologic Correlation, Genitourinary Radiology, 198 (3), 785–8. 12. Kaplan, I., Oldenburg, N.E., Meskell, P., et al. (2002) Real time MRI-ultrasound image guided stereotactic prostate biopsy, Magnetic Resonance Imaging, 20, 295–299. 13. Song, S.K., Qu, Z., Garabedian, E.M., Gordon, J.I., Milbrant, J. and Ackerman, J.J.H. (2002) Improved Magnetic Resonance Imaging Detection of Prostate Cancer in a Transgenic Mouse Model, Cancer Research, 62, 1555–8. 14. Yuen, J.S.P., Thng, C.H., Tan, P.H., et al. (2004) Endorectal Magnetic Resonance Imaging and Spectroscopy for the Detection of Tumor Foci in Men with Prior Negative Transrectal Ultrasound Prostate Biopsy, J. Urology, 171, 1482–68. 15. Nag, S., Beyer, D., Friedland, J., et al. (1999) American Brachytherapy Society (ABS) Recommendations for Transperineal Permanent Brachytherapy of Prostate Cancer, Int. J. Radiation Oncology Bio. Phys, 44 (4), 789–99. 16. Koseki, Y., Kikinis, R., Jolesz, F. and Chinzei, K., (2004) Precise Evaluation of Positioning Repeatability of MR-Compatible Manipulator Inside MRI, Medical Image Computing and Computer-Assisted Intervention, Saint-Malo, France, 26–30 September 2004, 192–9. 17. Koseki, Y., Koyachi, N., Arai, T. and Chinzei, K. (2003) Remote Actuation Mechanism for MR-compatible Manipulator Using Leverage and Parallelogram IEEE International Conference on Robotics and Automation, Taipei, Taiwan, 12–17 May 2003, 652–7. 18. Roebuck, J. PhD, MD, Brigham and Women’s Hospital, Private Communications, 2005–2007. 19. Wendt, O., Oellinger, J., Luth, T.C., et al. (2000) The effects of the use of piezoelectric motors in a 1.5-Tesla high-field magnetic resonance imaging system (MRI), Biomed Tech, 45, 20–5. 20. DiMaio, S.P., Fisher, G.S., Haker, S.J., et al. (2006) A Sytem for MRI-guided Prostate Interventions Proceedings of the International Conference on Biomedical Robotics and Biomechatronics, Pisa, Italy, 20–22 Febuary 2006, 68–73. 21. Taillant, E., Avila-Vilchis, J.C., Allegrini, C., et al. CT and MR Compatible Light Puncture Robot: Architectural Design and First Experiment, Medical Image Computing and ComputerAssisted Intervention, Saint-Malo, France, 26–30 September 2004, 145–52, 2004. 22. Stoianovici, D., Patriciu, A., Petrisor, D., et al. (2007) A New Type of Motor: Pneumatic Step Motor, IEEE/ASME Trans. on Mechatronics, 12, 98–106. 23. Fu, L., Ng, W.S., Liu, H., et al. (2005) Bouquet brachytherapy: Feasibility and optimization of conically spaced implants, Am. Brachytherapy Soc., 4, 59–63. 24. Devita, L.M., An MRI Compatible Manipulator for Prostate Cancer Detection and Treatment, M.S. Thesis, Massachusetts Institute of Technology, 2007. 25. Podder, T.K., Sherman, J., Clark, D.P., et al. (2005) Evaluation of Robotic Needle Insertion in Conjunction with In Vivo Manual Insertion in the Operating Room, Proceedings of the IEEE International Workshop on Robots and Human Interactive Communication, 13–15 August 2005, 66–72.

23 A Braille Display System for the Visually Disabled Using a Polymer Based Soft Actuator Hyouk Ryeol Choi1, Ig Mo Koo1, Kwangmok Jung2, Se-gon Roh1, Ja Choon Koo1, Jae-do Nam1 and Young Kwan Lee1 1

School of Mechanical Engineering, Chemical Engineering, Polymer System Engineering, Sungkyunkwan University, Korea 2 Pohang Institute of Intelligent Robotics, Korea

23.1 Introduction Tactile sensation is the most widely spread sensory function in the human body, so it is an essential part of the human perception mechanism. Complete or successful realization of the tactile function may place a cornerstone in the fields of robotics, virtual reality, medical engineering and so on. A tactile display is one of the most important communication devices and it conveys intellectual information in the form of skin stimulation. Although the graphical or visual device has been a typical form of the modern information transferring tool, the role of the tactile display has been extended to various applications, such as attaining reality in video games or guaranteeing accuracy in telemanipulation [1–4]. Especially, it has been a critical means for communication for the visually handicapped. Braille is a tool for exchanging information among the visually disabled and has been extensively used to transfer textual information. It consists of six pins arranged in pattern of a 3  2 matrix (a 4  2 matrix in the case of Chinese characters). Information is represented by stimulating human skin, usually fingertips, by vertical displacement of the pins. The tactile display device can be used as a refreshable

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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dynamic Braille. In particular, application of the display can also be expanded to a tablet capable of displaying textural or graphical information [5]. With this capability, even an entire web page can be delivered in a single display step. However, it is very difficult to enable Braille to deliver graphical information due mainly to the limitation of arranging massive Braille dots for high spatial density. Complicated and bulky driving mechanisms of conventional tactile displays hamper the development of a high resolution Braille-type tablet. According to a physiological study for standardization of Braille devices, the pin matrix density of a tactile display is typically up to 1 cell/mm2; actuating speed should be faster than 50 Hz; and energy density should be about 10 W/cm2 [6, 7]. Although the numbers are determined based on some experimental studies, the outcome of the display function is often deceptive since the sensitivity of the responses depends on the testing situation parameters such as speed, depth and strength of stimulation. Meanwhile, various mutated tactile display types have been introduced in order to accommodate human sensitivity, which normally varies from the fingertips to the palms. Many publications introduce several different types of tactile display devices that employ pneumatics, solenoids, voice coil, shape memory alloys, electrostatics or electroactive polymers [6–13]. Although previous developments deserve attention, most of them commonly suffer from low actuation speed due to complex actuation mechanisms. Furthermore, a complicated actuator design limits expansion to the tablet type application due to high manufacturing costs and low integration density. In this research, a new type of dynamic Braille display is presented. It employs a dielectric elastomer for the basis of the tactile display. It is constructed with a notably simple mechanical and electrical architecture. The proposed device is organized with a dual-layered array of tactile cells that generates vertical motion used to push up or down the Braille pins. These electrically driven tactile cells can generate either small-scale vibratory motion or linear displacement. They differ from conventional devices in softness and controllable compliance, cost effectiveness, simple manufacturability and high actuator density. Furthermore, the small size of the proposed concept enables the development of a high-density display device. In this chapter, in the Sections 23.2 and 23.3, the basic principle and design idea for the tactile cell is explained. The issues in the development of the Braille display device are discussed next (in Section 23.4) and then (in Section 23.5) exemplary feasible applications of the proposed tactile display are given.

23.2 Fundamentals of Actuation Principle The actuation principle of a dielectric elastomer actuator is similar to that of the electromechanical transduction of a two-plate capacitor (Figure 23.1). When a voltage is applied across the dielectric elastomer film coated with compliant electrodes on both sides, the elastomer is compressed in the thickness direction, while it is expanded in the lateral direction. Because of the contraction due to electrical charges stored in the surfaces, the actuation is generated. This physics couples the domains of mechanical and electrical energy to generate the energy transduction. The effective mechanical pressure along the thickness direction by a voltage input is given as:

A Braille Display System Using a Polymer Based Soft Actuator

V

V

Compliant Electrode

429

Dielectric Elastomer (a)

(b)

Figure 23.1 Actuation principle of a dielectric elastomer actuator: (a) voltage OFF; (b) voltage ON.

 ¼ "r "o E 2 ¼ "r "o



V t

2

ð23:1Þ

where E is the applied electric field, "o and "r are the electric permittivity of the free space and the relative permittivity, respectively. V represents the input voltage, and t denotes the thickness of the film. In other words, stress s along the thickness direction is proportional to the square of the applied electrical field. The introduction of high voltage across the dielectric elastomer film may result in a significant amount of strain in the thickness direction. Then, a relationship between initial thickness and strain in the direction is calculated as: t ¼ ð1 þ z Þt0

ð23:2Þ

where t0 and z are the initial thickness and the strain in the thickness direction, respectively. Combining Equations (23.1) and (23.2) and introducing an elastic modulus Y gives:  2 1 V "r "o z ¼ ¼ ¼ Y ð1 þ z Þt0 Y   z

 2 1 V 1 "r "o Y t0 ð1 þ z Þ2

ð23:3Þ

Then, a relationship between the input voltage, V, and the strain in thickness direction, z, is obtained by:  2 1 V "r "o 3z þ 22z þ z ¼ ð23:4Þ Y t0 The incompressibility condition gives:  ð1 þ  x Þ 1 þ  y ð1 þ  z Þ ¼ 1

ð23:5Þ

ð1 þ r Þ2 ð1 þ z Þ ¼ 1

ð23:6Þ

Since x ¼ y ¼ r , in the case of the circular elastomer film:

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where r represents the strain along the radial direction. Then, r can be rewritten as: 1 r ¼ pffiffiffiffiffiffiffiffiffiffiffiffi 1 þ z

1»

1 z 2

ð23:7Þ

Therefore, mechanical actuation can be acquired via pressure causing axial contraction or lateral expansion of the dielectric elastomer film. Further details are well elaborated in previous reports [15, 16].

23.3 Design of Tactile Display Device Depending on the design and fabrication technique, the dielectric elastomer actuator can be implemented in various forms and generate a wide range of motions from micro to macro [15, 16]. The stimulating cell proposed in this chapter can provide durable and robust actuation since it is configured to generate mechanical output orthogonal to the surface of the elastomer film without employing the pre-strain frequently used in the previous studies [14]. As shown in the schematic diagram in Figure 23.2, this design provides relatively large orthogonal displacement compared to the size of the cell and allows very easy fabrication. The operating principles of the proposed stimulating cell are described in detail below. Moving

a

b = a(1 + δa)

r

Frame

(a)

Figure 23.2 actuation.

h

Electrode A: + Electrode B: –

Dielectric elastomer

θ

(b)

Basic actuation modes of the stimulating cell: (a) without actuation; (b) with

An incompressible thin circular elastomer film is attached inside a rigid cylindrical boundary frame. When a voltage is applied across the dielectric elastomer film, it is compressed along the axial (or thickness) direction and, thus, it expands along the radial direction. The expansion along the radial (or lateral) direction causes concave or convex bending of the elastomer (which may be called ‘buckling’). For the typical construction, a simple geometric relation for the circumference of the convex b is calculated by: b ¼ að1 þ a Þ ¼ r

ð23:8Þ

where r is the radius of curvature,  is the corresponding angle and a is the initial (or not deformed) radius of the film. a is the induced strain in the radial direction, which is calculated from Equations (23.3) and (23.7). In addition:

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431

   a ¼ rsin 2 2

ð23:9Þ

 ¼ 2ð 1 þ  a Þ sinð=2Þ

ð23:10Þ

Then, Equation (23.8) becomes:

Applying the Taylor series expansion to the previous result gives:     sin » 2 2

ð=2Þ ð24 2 Þ ¼ 3! 48

ð23:11Þ

By substituting Equation (23.11) into Equation (23.10), an algebraic relation between  and a can be determined: sffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi  ffi 1  ¼ 24 1 ð23:12Þ ð1 þ  a Þ Thus, the trigonometric relationships produce the convex height, h, such as:     h ¼ r 1 cos 2

ð23:13Þ

where r is to be determined by Equation (23.9).

23.4 Braille Display System How a Braille cell operating as per the presented mechanism is fabricated and tested is described in this section. 23.4.1

Fabrication

Construction of the actuator is as follows. A thin cylinder shaped polymer membrane such as a coin is coated with carbon electrodes. Nominal thickness of the elastomer membrane, whish is made with the silicone KE441 by ShinEtsu, is about 50 mm. In order to obtain an actuation force large enough to be used for the tactile display, several laminated membrane actuators are stacked and combined with a rigid circumferential loop frame (Figure 23.3). The total thickness of the elastomer stack is about 750 mm. Note that the actual diameter of the elastomer stack is slightly larger than the rigid circumferential frame. Introduction of this oversized fitting design produces predeformation of a convex that guarantees unidirectional actuation mode. The prototype fabrication of the actuator and its cross-sectional cut are shown in Figures 23.4a and 23.4b, respectively.

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Dielectric Elastomer

Non-electric area Frame

Electrodes Dielectric Elastomer dr d df

Figure 23.3 Construction of the proposed actuator.

50 µm

Dielectric elastomer

Carbon electrode

(a)

(b)

Figure 23.4 Prototype of the actuator: (a) actuator unit; (b) cross-section view.

23.4.2

System Outline

A typical Braille display unit is constructed with six stimulating pins that are arranged in a 3  2 array format. An array normally represents a character as defined by the Braille alphabet. The standard Braille display unit is illustrated in Figure 23.5. In the present work, a Braille display unit is constructed with the introduced tactile cells arranged in the format defined by the standard Braille display. The construction concept is depicted in Figure 23.6. Although the dielectric elastomer based tactile cell is driven with high voltage electricity, users have no direct contact to the actuator surface. A Braille pin made from insulating material is the only contact to human fingertips. In addition to the pre-deformed convex feature of a cell, note that the directional ball is placed underneath each cell in order to guarantee unidirectional actuation. Packaging the six pin actuators and corresponding electric wires in a constrained small space might

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2.4

φ1.6 4

2

5

3

6

2.4

8

2.4

1

5.6

0.8

∗ unit : mm

12

Figure 23.5

Standard Braille cell consisting of six dots.

Braille pin Upper frame 1

Actuator 3 Actuator 2

Actuator 1 Direction ball

Silicone Lower frame 1

Upper frame 2

Actuator 6 Actuator 5

Actuator

Direction ball

Silicone Lower frame 2

Gnd 2

Gnd 1

Figure 23.6

Vcc 1

Vcc 3 Vcc 2

Exploded view of proposed Braille display.

require an expensive manufacturing process. For this reason, a dual layer construction is introduced in order to alleviate the fabrication problems caused by the high-density small apparatuses. By allocating three pins in each layer with a staggered pattern, interferences caused by complicate wiring can be minimized. Each layer is shown in Figure 23.7. As shown in Figure 23.8, the height of the fully assembled device is approximately 9 mm excluding the length of the terminals. Since each Braille cell is to be fully modularized for convenient installation, each unit can be plugged onto a circuit board with ease. With this

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Biomedical Applications of Electroactive Polymer Actuators

Hole Actuator cell

Upper layer

Figure 23.7

Lower layer

Top and bottom layers of a Braille cell.

Figure 23.8 Assembled Braille cell: (a) assembled Braille cell; (b) side view; (c) top view.

simple drop-in feature, a number of Braille cells can easily be combined so that a Braille tablet may be manufactured by arranging many Braille cells in a matrix format as illustrated in Figure 23.9. 23.4.3

Experiments

To evaluate the performance of the proposed Braille display, psychophysical tests have been performed using visually impaired persons. Seven individuals participated in the test, six of them were males and one was a female with an overall average age of 32.6 years. All of them were employees of the Korean Association of the Visually Impaired. One of them had weak vision and the others were completely blind. The software environment graphical

435

17 mm

A Braille Display System Using a Polymer Based Soft Actuator

18.8 mm

Figure 23.9

Braille tablet made by assembling six modular Braille cells.

user interface (GUI) used for the testing is shown in Figure 23.10; output voltage, frequency and duration time can be adjusted by a computer so that the Braille information can be encrypted and delivered to the display unit. Although only numeric characters were used for the testing, nothing is precluded for the implementation of any characters defined by Braille.

Figure 23.10

User interface.

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Figure 23.11 Psychophysical experiment of the proposed Braille device.

As shown in Figure 23.11, two adjacent Braille cells were read by the subjects and the data for two types of recognition rates – Hit Recognition Rate (HRR) and Number Recognition Rate (NRR) – were obtained. HRR denotes the rate at which a subject recognizes movement of the Braille dots and NRR represents whether a subject reads the character correctly as the Braille dots are activated. In the experiments, the HRR and NRR of each subject have been tested when the actuating frequency of the Braille pins was at 15 Hz, which is the normal read out speed of Braille readers. The results of the tests are shown in Table 23.1. In the experiment, the HRR reaches up to about 80 % and the NRR shows a maximum of 41 %. The test results of 80 % for HRR and 40 % for NRR are much better than originally expected.

Table 23.1 Subject No. 1 2 3 4 5 6 7

Psychophysical test results – HRR and NRR HRR (%)

NRR (%)

69 51 54 60 77 73 63

30 21 19 29 40 41 32

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437

A main concern before conducting the experiment was the fact that the subjects did not have any experience on the proposed Braille display device. Realizing the sharp sensitivity of blind people, the recognition rates are sure to improve provided that they have more exposure to the device. In addition, it should be emphasized that the recognition rates are greatly affected by two major factors, the regularity of the Braille pin’s height and the activating frequencies; the Braille unit used in this test was hand-crafted and the consistency was not tightly controlled. The higher driving frequency will, of course, aggravate the recognition rate. Although the maximum tactile sensitivity of a human fingertip is reported to be about 250 Hz according to a physiological study [17], the subjects that participated in the test complained about weakening tactile sensation when the tests were driven at high frequencies up to 60 Hz. Unlike a static Braille characters, the proposed dynamic Braille device might have generated a small-scale vibration from the Braille pins, confusing the subjects and adversely affecting the recognition rates. However, due to the intrinsic low-pass characteristics of polymeric materials such as dielectric elastomers, there is no situation that a high frequency vibration caused by high frequency noise can be combined with the Braille signal. This problem should be further investigated with improved device fabrication and a larger subject size. Nonetheless, the presented device proves its industrial feasibility from the experiments presented. Furthermore, it was discovered from the experiment that maintaining the regularity of the Braille pin’s height might be one of the major factors for precise recognition. Even a small variation of the height was easily detected by the subjects and had a tendency to confuse them. The problem could be eliminated if the fabrication was supported by sophisticated machines. Besides, it is likely that the number of subjects should be increased for the precise evaluation of the proposed device.

23.5 Advanced Applications In this section, exemplary feasible applications of the proposed tactile display are given. 23.5.1

Wearable Tactile Display System

A new type of tactile display device, that is a wearable tactile display device, has an array of multiple tactile stimulating cells embedded on a flexible polymer substrate (Figure 23.12). The immediate outcome of the present research is shown in Figure 23.13. The dielectric elastomer actuator provides a number of benefits for the actuator construction, namely excellent flexibility with comfortable softness, ease of fabrication, miniaturization and cost effectiveness [15, 16]. Attaining the flexibility, the device can be adapted to various geometric configurations and can be worn on any part of the human body, such as fingertip, palm or arm. The device is actually a polymer sheet that has an embossed array of soft actuators. The polymer sheet is preformed with an embossed pattern in order to secure the moving direction of the actuator. The embossed pattern determines the direction of buckling and, thus, the arbitrariness of the direction of movement can be removed. Since the construction of the device is simple as described, the flexibility of the polymer substrate is preserved, and the fabrication process is extremely easy to implement in reality. The distal part of a human finger is not only one of the most sensitive parts of the human body,

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Biomedical Applications of Electroactive Polymer Actuators

Figure 23.12 Proposed tactile cell: (top) exploded layer and cross-sectional view of tactile simulating cell; (bottom) assembly picture of tactile stimulating cell.

Figure 23.13

Illustration of the proposed wearable tactile display and flexible characteristics.

but also extensively used for the detection of environments. Sufficient flexibility, lightweight and simple ‘one-shot’ fabrication of the proposed actuator enable the miniaturization of the tactile display device without the assistance of any complicated electromechanical parts. The schematic diagram provided in Figure 23.14 shows the major parts of the device. The functions of each part are:

A Braille Display System Using a Polymer Based Soft Actuator

Figure 23.14

439

Exploded view of the proposed wearable tactile display.

 Frame Film: 50 mm thick elastomer film. It is not covered with electrodes and works as a constraint to the stimulating cell.

 Actuator Film: Elastomer film with a thickness of 200 mm. It is constructed by stacking

eight layers of dielectric elastomer and contains 4  5 arrayed tactile stimulating cells. Each layer is 25 mm in thickness.  Protective Film: Dielectric elastomer film with 5 mm thickness. It keeps the human skin from direct contact with tactile stimulating cells. Rolling up the presented actuator unit to a thimble shape tube, as shown in Figure 23.15, the device can easily fit on the human fingertip. The total active area for the device is 11  14 mm, and the centers of the tactile stimulating elements are apart 3 mm. Each element is 2 mm in diameter and the initial convex height is 0.1 mm. The commercially available silicone (CF19-2186, Nusil Company) is used for the fabrication of all parts. The entire device is very flexible and light like a bandage enough to be worn at the fingertip. To supply the electrical energy, a flexible and thin enamel wire (0.18 mm thickness) is used, which is helpful to prevent interferences with motions of the human finger. The system consists of a high-voltage source, switching circuits and an embedded controller with universal serial bus (USB) communication interface.

Figure 23.15 Assembled 3-D model of the wearable tactile display: (top) 3-D view of assembled display; (bottom) cross-sectional view.

440

23.5.2

Biomedical Applications of Electroactive Polymer Actuators

Virtual Reality Tactile Display

Combining the proposed actuator unit to a virtual keyboard, a VRTD (Virtual Reality Tactile Display) system that transfers more realistic keyboard touch feeling to the user can be constructed. The compact version of the commercially available virtual keyboards provides the advantage of high mobility. However, they sometimes cause confusion and discomfort to users since the users have to hit the hard ground. By wearing the proposed tactile display device on the end of fingertips, the level of keyboard touch feeling can be significantly improved. As soon as the user hits a character, the tactile device will generate keyboard touch feeling. The proposed VRTD system configuration and concept are illustrated in Figure 23.16. A virtual keyboard (I-tech Co.) and a user interface (Visual studio.Net based on Windows XP) used for controlling tactile display are shown in Figure 23.17. The functions of the GUI are:

 USB Connect: provides the connection of the PC to the micro controller.  Real time keyboard scan: presents keyboard characters by real time scanning virtual keyboard push buttons.

 Tactile display (Braille code): inputs characters for virtual keyboard and converts to the    

corresponding Braille codes. Also it displays the Braille code on the tactile display. Long word: sends a long sentence as a sequential word. Output control: controls displacement of the actuator by modulating applied voltage. Frequency: controls the actuation speed to provide varying tactile feelings. Display time: controls displaying duration of each character.

Figure 23.16

Virtual reality tactile display system.

A Braille Display System Using a Polymer Based Soft Actuator

441

Figure 23.17 Experiments with the proposed virtual keyboard.

Recognizing the working mechanism of the proposed VRTD, a visually impaired user also can benefit from the device. The device can deliver the input characters to the tip of finger if coded as the Braille rule so that the user can read out his input as soon as the characters are typed in.

23.6 Conclusions In this chapter, a Braille display device fabricated with polymer actuators is proposed and tested. The presented device offers significant advantages over the existing display methods. The device provides notable flexibility, simple manufacturing and cost-effectiveness. In addition, it has many advantages over the existing electromechanical counterparts when it needs to be used for the tactile display. On the other hand, the high driving voltage requirement and low bandwidth might be some of the drawbacks. They might hamper the application of the device to sophisticated tasks. Since the proposed device is organized with ‘soft’ actuators, it might assist to alleviate psychological or physical disturbances that could be accumulated during the lengthy operation of Braille devices by the blind.

References 1. Johansson, R. S. and Vallbo, A. B. (1983) Tactile Sensory Coding in the Glabrous Skin of the Human Hand, Trends in Neurosciences(TINS), 6, 1, 27–32. 2. Hannaford, B., Wood, L., McAfee, D. and Zak, H. (1991) Performance Evaluation of a Six Axis Generalized Force Reflecting Teleoperator, IEEE Trans. Systems, Man, and Cybernetics, 21, 620–3.

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3. Yoshikawa, T., Yokokohji, Y., Matsumoto, T. and Zheng, X. (1995) Display of Feel for the Manipulation of Dynamic Virtual Objects, ASME J. Dynamic Systems, Measurement and Control, 117, 4, 554–8. 4. Moy, G., Wagner, C. and Fearing, R. S. (2000) A compliant tactile display for teletaction, Proceedings of the IEEE International Conference on Robotics and Automation, 3409–15. 5. Shinohara, M. and Shimizu, Y. (1998) Three-Dimensional Tactile Display for the Blind, IEEE Trans. Rehabilitation Eng., 6, 3, 249–56. 6. Jungmann, M. and Schlaak, H. F. (2002) Taktiles Display mit elektrostatischen Polymer Aktoren, Proceedings Internationales Wissenschaftliches Kolloquium(in German). 7. Asamura, N., Shinohara, T., Tojo, Y., et al. (2001) Necessary Spatial Resolution for Realistic Tactile Feeling Display, Proceedings of the International Conference on Robotics and Automation, 1851–6. 8. Caldwell, D. G., Tsagarakis, N. and Giesler, C. (1999) An Integrated Tactile/Shear Feedback Array for Stimulation of Finger Mechanoreceptor, Proceedings of the IEEE International Conference on Robotics and Automation, 287–92. 9. Debus, T., Jang, T. J., Dupont, P. and Howe, R. (2002) Multi-channel Vibrotactile Display for Teleoperated Assembly, Proceedings of the International Conference on Robotics and Automation, 592–597. 10. Tang, H. and Beebe, D. J. (1998) A Microfabricated Electrostatic Haptic Display for Persons with Visual Impairments, IEEE Trans. Rehab. Eng., 6, 5, 241–8. 11. Spinks, G. M., Wallace, G. G., Ding, J., et al. (2003) Ionic Liquids and Polypyrrole Helix Tubes: Bringing the Electronic Braille Screen Closer to Reality, Proceedings of the SPIE 10th Annual Symposium on Smart Structures and Materials: Electroactive Polymer Actuators and Devices (EAPAD), 372–80. 12. Konyo, M., Akazawa, K., Tadokoro, S. and Takamori, T. (2003) Wearable Haptic Interface Using ICPF Actuators for Tactile Feel Display in Response to Hand Movements, J. Robotics and Mechatronics, 15, 2, 219–26. 13. Taylor, P. M., Hosseini-Sianaki, A. and Varley, C.J. (1996) An Electrorheological Fluid–based Tactile Array for Virtual Environments, Proceedings of the IEEE International Conference on Robotics an Automation, 18–23. 14. Pelrine, R., Kornbluh, R., Pei, Q. and Joseph, J. (2000) High-Speed Electrically Actuated Elastomers with Over 300 % Strain, Science, 287, 836–839. 15. Choi, H. R., Lee, S. W., Jung, K. M., et al. (2004) Tactile Display as a Braille Display for the Visually Disabled, Proceedings of the IEEE/RSJ International Conference on Intelligent Robotic Systems, September/October 2004, 2, 1985–90. 16. Koo, I. M., Jung, K. M., Koo, J. C., et al. (2008) Development of Soft Actuator Based Wearable Tactile Display, IEEE Trans. Robotics, 4, 3, 549–59. 17. Kaczmarek, K. A. (1991) Electrotactile and Vibrotactile Displays for Sensory Substitution Systems, IEEE Trans. Biomed. Eng., 38, 1, 1–6.

24 Dynamic Splint-Like Hand Orthosis for Finger Rehabilitation Federico Carpi, Andrea Mannini and Danilo De Rossi Interdepartmental Research Centre ‘‘E. Piaggio’’, School of Engineering, University of Pisa, Italy

24.1 Introduction This chapter presents preliminary investigations on the use of dielectric elastomer actuators as active components of a specific type of orthotic systems, known as ‘hand splints’, used for finger or hand rehabilitation. In general, an orthotic system is a structure applied externally to a part of the body, in order to either correct deformity, improve function or relieve symptoms of a disease. Accordingly, an orthotic system provides a means to support, assist or restore neuromuscular functionalities of a body part. Hand splints have today well established orthopedic applications as orthotic systems to immobilize either the entire hand or just one or more fingers. They can be adopted for different purposes, including post-surgical or post-trauma immobilizations and articulation-corrective or articulation-supporting actions. Such typical examples of uses refer to so-called ‘static hand splints’. More generally, hand splints can be divided according to the scheme reported in Figure 24.1. This chapter is focused on ‘dynamic hand splints’, whose state of the art is briefly presented below.

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Figure 24.1

A possible classification of hand splints.

24.2 Passive Dynamic Hand Splints: State of the Art Dynamic hand splints consist of mechanically compliant orthoses conventionally used to perform rehabilitation of fingers or the overall hand in correspondence of specific allowed motions. In particular, so-called ‘passive dynamic hand splints’ frequently include elastic bands or springs, which merely exert a passive resistance to voluntary elongations of one or more fingers. Figures 24.2a to 24.2c show some examples. These kinds of orthotic systems enable rehabilitation of fingers that still can be moved voluntarily against the recovery force of the counteracting elastic component. Commercial splints used for such purposes generate maximum forces typically of the order of 1–10 N.

Figure 24.2 Examples of passive dynamic hand splints equipped with different passive components: elastic bands ((a) Phoenix outrigger, adapted from [1], (b) LMB Wrist Extension Assist, adapted from [2]); linear springs ((c) Rolyan adjustable outrigger, adapted from [3]); and torsional springs ((d) DeROM Dynamic Range of Motion Wrist Splint, adapted from [4]).

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Although such an approach represents a very simple and useful means to implement elementary rehabilitation exercises, it is intrinsically affected by an evident limitation: the impossibility of modulating the counteracting action. In fact, the mechanical compliance of the system is pre-defined by the fixed mechanical properties of the adopted elastic component. In contrast, the possibility of exploiting a variable compliance could significantly enhance the rehabilitation versatility and efficacy; in fact, it would permit exercise of the finger against variable reaction forces, to be selected according to desired rehabilitation plans. The variation of the system compliance can be obtained in state-of-the-art passive dynamic hand splints by using either interchangeable elastic components (Figures 24.2a to 24.2c) or adjustable torsional springs (Figure 24.2d); these have to be respectively replaced or adjusted every time it becomes necessary. In order to avoid this, so-called ‘active dynamic hand splints’ represent an attractive alternative.

24.3 Active Dynamic Hand Splints: State of the Art The purely manual regulations necessary for passive splints provide, of course, intrinsic limitations with respect to more useful electrical regulations. Such functionality is offered by ‘active dynamic hand splints’. They are conceived as orthotic systems with embedded electromechanical actuation mechanisms, aimed at enabling an electrical control of the most relevant mechanical features of the system. In particular, active splints are aimed at delivering rehabilitation forces that are precisely and continuously controllable by means of an electrical input, even in closed-loop systems. Accordingly, the patient can undergo accurate rehabilitation programmes by means of a useful control of the most significant rehabilitation parameters. These could consist, for instance, in the fundamental features of a force signal, such as its waveform, amplitude, frequency, duration, rise time, repetition rate and so on. These parameters could be either selectable among predefined sets or even continuously adjustable during the training phase, according to the specific response of the patient. In order to develop such active orthoses, the choice of the actuation technology is a key issue for the efficacy of the resulting system. Several electromechanical transduction mechanisms to endow hand systems (not necessarily a splint) with actuation functions have been studied so far. These technologies include pneumatic devices [5–7] (Figures 24.3a and 24.3b), electromagnetic motors [7, 8] (Figure 24.3c) and shape memory alloys [9] (Figure 24.3d). Additionally, useful means to electrically modulate resistive forces for hand rehabilitation devices have been demonstrated recently by using electrorheological fluids [10]. Despite the attractive performances offered by all these actuation technologies, such approaches are typically characterized by heavy, stiff and bulky devices. They complicate the structure of the orthotic system and make it uncomfortable and less easily wearable and portable. In particular, pneumatic actuators (e.g. McKibben-like devices) can generate large pressures with an output mechanical characteristic that mimics nicely that of natural muscles; nevertheless, they require gas compressors and regulators and show a relatively low versatility. Shape memory alloys (e.g. Nitinol) are advantageously available as wires and can exhibit large stresses and high energy densities; however, they typically show

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Figure 24.3 Examples of active hand systems equipped with different types of actuators: pneumatic devices ((a) adapted from [6] Copyright (2004) IEEE, (b) Rutgers MII glove, adapted from [7] Copyright (2002) IEEE); electromagnetic motors ((c) Cyber grasp/Cyber glove, adapted from [7] Copyright (2002) IEEE) and shape memory alloy wires ((d) adapted from [9]).

relatively long response times, low efficiency and hysteresis. Electromagnetic DC motors, the most used in biorobotics, are available in compact sizes with attractive performances; nevertheless, they offer the maximum output power at a maximum input voltage, which means maximum angular speed (at least with no load). Therefore, DC motors typically require adequate transmissions (gears) to reduce the angular speed and increase the torque; this necessarily complicates the structure and increases its encumbrance. Furthermore, DC motors are heavy, noisy and stiff, so that they have to be also combined with springs to make the system mechanically compliant. Therefore, new types of actuation technologies are today being searched for to develop more viable active orthotic systems.

24.4 Proposed Concept: Dynamic Splint Equipped with Dielectric Elastomer Actuators Within such a context, dielectric elastomer (DE) actuators may be regarded as an interesting opportunity, deserving detailed investigations. In fact, their intrinsic material properties and working principles offer unparalleled features [11–13]. These are potentially suitable to provide the orthotic system with adequate functional properties, while making it, at the same time, at least lighter, more flexible and comfortable and, accordingly, more easily wearable and portable. This chapter describes an actuated dynamic hand splint with variable compliance regulated by new DE contractile actuators. The concept is described below, with reference to Figure 24.4.

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Figure 24.4 Dynamic hand splints: (top) example of a passive dynamic splint equipped with an elastic band; (bottom) schematic drawing of the proposed concept: the passive elastic band is substituted with active elastic actuators.

To develop a wearable dynamic hand splint with electrically variable compliance, the concept described here relies on the use of linear contractile DE actuators. The idea is to use these actuators as active elastic substitutes of the linear passive elastic components of traditional splints, as sketched in Figure 24.4. In order to easily assemble a linear contractile DE actuator, a recently described simple structural configuration, known as a ‘folded actuator’ [14], can be advantageously adopted. Such a device consists of a monolithic strip of dielectric elastomer, which is first coated with compliant electrodes and then folded several times in order to obtain a compact structure (Figure 24.5). Following the application of a high voltage difference between the compliant electrodes, the actuator operates an electromechanical transduction resulting in a compressive stress that tends to squeeze the thickness of the insulating layer. Accordingly, an axial contraction combined with lateral expansion can be achieved if the structure is actually allowed to shorten [14]. For any given driving electric field, E (i.e. the ratio between the applied voltage, V, and the thickness, d, of the elastomeric layer), the generated active stress (force per unit cross-section), p, depends on fundamental electric properties of the elastomer, according to the expression [11, 12]: p ¼ "0 "r E2 ¼ "0 "r ðV=dÞ2

ð24:1Þ

where "r is the relative dielectric constant of the elastomer and E0 is the dielectric permittivity of vacuum. This compressive stress (pressure) squeezes the elastomer and determines its actuation performance.

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Figure 24.5 Dielectric elastomer folded actuator: picture of a prototype sample and schematic drawing of the device structure. An electrical activation generates a compressive force along the axial direction.

Linear contractile elastic actuators readily appear as a particularly suitable choice for the application considered here. By equipping a splint with such actuators, their electrical activation can be used to vary the compliance of the system; this is aimed at modulating the antagonistic force that reacts to voluntary finger movements, according to programmable rehabilitation exercises. In order to implement this concept, the splint was conceived as a wearable system equipped with parallel folded actuators, connected to the pulling finger through an inextensible wire (Figure 24.6). 2 pulley wire H

1

actuator

+ –

F1 1

Frehab, 2

F2 P1

splint α

o

ϑ P2

2

Figure 24.6 Schematic drawing of a finger connected to dielectric elastomer contractile actuators arranged aboard a splint. Position 1: the finger axis is aligned with the splint; Position 2: the finger axis is voluntarily rotated by the patient of an angle a around a fulcrum O.

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24.5 Splint Mechanics To obtain a simple analytical description of the basic mechanics (in static condition) that are behind this application, consider a splint equipped with N parallel actuators (where N has to be defined according to the problem constraints, as specified in the following). Figure 24.6 presents a schematization of the system, by showing a rotation a of a finger around an axis crossing a fulcrum O. Any voluntary flexion of the finger lengthens the actuators. Each of them offers a counteracting force, whose modulus F is the sum of a passive component (elastic recovery) and an active component (electrically induced). Simple calculations adopting a linear elastic model for the elastomer (in the hypothesis of sufficiently small strains, e.g. up to 10%), lead to the following total force generated by the N parallel actuators:   A0 L0 L  L0 " 0 "r V 2 FffiN Y ð24:2Þ þ 2 L L0 d0 ð1þðL  L0 Þ=L 0 Þ 2 where A0 is the rest value of the cross-section of each actuator, L0 and L are, respectively, the rest and the actual values of its length, d0 is the rest thickness of the dielectric layer and Y is the Young’s modulus of the elastomer. Equation (24.2) shows that, as the finger flexion progressively lengthens the actuators, the active component of F decreases as the inverse of the square lengthening. In order to fully exploit the actuation, such a decrease should be minimized. This can be achieved, for any given finger flexion, that is any given actuator lengthening L  L0, by maximizing L0. The effective force useful for rehabilitation purposes is the component of F that acts perpendicularly to the finger, since the tangential component provides just a shear stress on the soft tissues. To maximize this component, the splint can be equipped with a support able to guide, eventually through a pulley, the pulling wire perpendicularly to the finger, as shown in Figure 24.6. Making the position of the pulley adjustable along the support can compensate possible geometrical variations for individuals with different hand sizes. The mechanical compliance, C, of the considered system can be defined as coincident with that of the N parallel actuators, that is the ratio between the actuator lengthening and the voltage dependent applied force: C¼

L  L0 FðVÞ

ð24:3Þ

The aim of this chapter is to show how C can be electrically modulated, for any given L  L0, through a variation of F according to a regulation of the driving voltage V.

24.6 Dimensioning of the Actuators The technical design of the splint and the related dimensioning of the actuators have necessarily to take into account, as a fundamental specification, the maximum antagonist force that the actuators should be able to offer. An analysis of the state-of-the-art of hand splints for finger rehabilitations showed a chronic lack of extensive and comparable data on suitable rehabilitation forces adopted in the clinical practice. However, as a special case,

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conventional passive splints equipped with elastic bands adopted for post-surgical hand rehabilitations were found to work in a range which frequently extends up to a few Newtons [15, 16]. Such a working range was regarded here as the target of this first study on dynamic splints based on DE actuators. To satisfy this requirement, both the number of actuators to be used, their size, their constituent material and the performances achievable from each of them should be considered. This work adopted a ‘standard’ type of prototype folded actuator currently being fabricated in our laboratory; these are characterized by d0 ¼ 1 mm and A0 ¼ 16  21 mm2. According to their typical performances [14], N ¼ 3 parallel devices were used. The structure of the developed splint (reported in the next section) allowed a maximum length for the actuators of 96 mm, while a maximum flexion corresponding to (L  L0)max ¼ 10 mm was considered; accordingly, L0 ¼ 85 mm was used.

24.7 Prototype Splint A prototype splint was designed so as to be able to house aboard it the following components: the three parallel actuators required, the high voltage (HV) electronics to drive them, a load cell to continuously monitor the force exerted by the actuators and an aluminium rod supporting a pulley, used to guide a nylon wire which connected the actuators to the pulling finger. A CAD design of the system is shown in Figure 24.7.

Figure 24.7 Design of a hand splint equipped with linear dielectric elastomer actuators, their high voltage electronics and a load cell.

As a support for the pulley, a frame able to slide along the rod was considered, in order to make its position easily adjustable for patients with different hand sizes. The actuators were made of a silicone elastomer and were fabricated with the same materials and procedure described in [14].

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The HV driving of the actuators was accomplished using three miniature DC to HV DC converters (Q101-5, EMCO High Voltage Corporation, USA) able to provide an output voltage of 0–10 kV for an input of 0–5 V [17]. A picture of this type of HV module is given in Figure 24.8. The overall electronics was packaged in a plastic box, arranged on the rear of the splint (Figure 24.7).

Figure 24.8 Miniature high voltage power supply manufactured by EMCO Ó (Reproduced from www.emcohighvoltage.com [17]).

A plastic prototype splint developed according to the above design is shown in Figure 24.9. This prototype had a total weight (not optimized, of course) of 530 g. The overall system was designed to be connected to a laptop computer through a data acquisition board, controlled by means of custom software. A user-friendly graphic interface (Figure 24.10) allows the waveform, amplitude, frequency and duty cycle of the actuator driving voltage to be set easily; signals of both the supplied voltage and the continuously monitored force are visualized, so as to allow the user to quantify and evaluate the efficacy of the rehabilitation exercise.

24.8 Performance of the Prototype Splint Results of the characterization of the prototype splint are presented in Figure 24.11, which reports the total force to be applied by the finger in order to achieve different amplitudes of actuator lengthening L  L0, at different driving voltages. For any given lengthening, each plot shows a bias (corresponding to the passive component) plus a voltage-dependent quadratic contribute (active component), as foreseen by Equation (24.2). As an observation, the curve for a null lengthening shows a bias force (i.e. the purely passive component, for 0 kV) which is not null. This is simply due to a slight pre-tensioning applied to the actuators in order to keep them approximately straight (the length

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Figure 24.9 Prototype dynamic hand splint equipped with silicone folded actuators, high voltage electronics and load cell.

corresponding to such a pre-loading was assumed as the new L0, from which all the variations were calculated). Force and lengthening data were used to calculate the compliance of the system (Equation (24.3)), whose plot as a function of the applied voltage is reported in Figure 24.12. These results demonstrate that, as expected, the use of the actuators actually enables electrical modulations of the system compliance. This represents a proof of concept of the intended application, showing the benefit of equipping the splint with electrically controllable elastic components over purely passive elastic components. This permits the need for manual regulation, or even substitution of the resistive components, to vary the rehabilitation exercise to be avoided. Nevertheless, these data show also that the achievable active modulation of the force was rather moderate in this first prototype system; in fact, it allowed a range of 1 N. Accordingly, although these data provide the first proof of concept of the proposed technique, they also suggest a need of further improvements. In particular, an increase of the electromechanical performance of the prototype actuators considered in this work is required. This could be achieved be means of three parallel approaches, whose

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Figure 24.10 Graphic interface of the control software developed for the hand splint. 9 L – L0 = 6 mm (7.1% L0)

8

Total force F [N]

7 L – L0 = 4 mm (4.7% L0)

6 5

L – L0 = 2 mm (2.3% L0)

4 3

L – L0 = 0 mm (0% L0)

2 1 0

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2

3 Voltage V [kV]

4

5

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Figure 24.11 Force–voltage plots recorded from the prototype splint for different amplitudes of actuator lengthening.

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Compliance [mm/N]

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0.78 0.76 0.74 0.72 0.70 0.68 0.66 0.64 0.62 0.60 0.58 0.56 0.54 0.52 0.50

L – L0 = 6 mm (7.1% L0)

L – L0 = 4 mm (4.7% L0)

L – L0 = 2 mm (2.3% L0)

0

Figure 24.12

1

2

3 Voltage V [kV]

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6

Compliance–voltage plots for different amplitudes of actuator lengthening.

implementation imply considerably different times and challenges. These approaches are briefly listed here, from the easiest to the most challenging: (i) a readily feasible adoption of dielectric elastomers with a higher dielectric strength, so as to extend the upper limit of applicable electric fields, increasing the maximum achievable stress, accordingly Equation (24.1); several commercial silicones, for instance, could be easily employed for such a purpose; (ii) development of improved fabrication processes, specifically studied to allow handling of elastomer layers with reduced thickness, so as to lower the driving voltages Equation (24.1); and (iii) development of new dielectric elastomers exhibiting superior intrinsic electromechanical properties, namely a higher dielectric constant, so as to reduce the driving electric field (and, thus, the driving voltage) required to achieve any definite stress (Equation (24.1)). Fulfilment of such needs is currently being pursued [18–20].

24.9 Future Developments Beyond the necessary improvement of performance described above, application-orientated developments currently being considered are focused on two types of specific aims, described in the following sections. 24.9.1

Magnetic Resonance Imaging-Compatible Hand Splint

A new version of the hand splint is currently being developed in order to obtain a prototype compatible with the environment of Magnetic Resonance Imaging (MRI). An MRIcompatible hand splint could be used to perform rehabilitation exercises within an MRI scanner, so as to allow functional evaluations of the rehabilitation efficacy.

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MRI and, in particular, so-called functional MRI (fMRI) is used today as a powerful tool for monitoring progress of rehabilitation therapies. This type of application requires MRIcompatible rehabilitation tools to be introduced inside (or very close to) the MRI scanner. Accordingly, in order to drive such systems, MRI-compatible actuators are needed. In general, the two fundamental requisites that define the MRI-compatibility of a device are: (i) its capability of withstanding the strong magnetic fields of the MRI scanner and sensitive imaging sequences, without any significant performance degradation; and (ii) its capability of functioning without inducing any significant artifacts on the acquired images. Conventional actuation technologies typically are not or are scarcely suited for this field of application [21]. In fact, electromagnetic motors are intrinsically not applicable, due to their working principle. Hydraulic and pneumatic actuators can exhibit MRI-compatible properties, although they are usually bulky and heavy. As an exception, piezoelectric actuators combine a good MRI compatibility with a compact structure, so that they currently represent the preferred choice for MRI-compatible actuation technologies. Nevertheless, they still exhibit some problems, including possible electromagnetic interference due to AC driving required by the working principle (as in the case of ultrasonic rotary motors), very small intrinsic strains (up to 0.1 %) that require mechanical amplification (with complication of the structure and encumbrance of the overall system) and relatively high costs. Therefore, new actuation materials and technologies are of interest, to develop new MRI-compatible mechatronic systems capable of extended and improved capabilities (at least for complementary needs). Dielectric elastomer technology could represent a useful and valuable possibility, enabling the development of systems more easily wearable (thanks to a mechanical flexibility, low specific weight and small volume), with no acoustic noise (so as to improve the patient’s comfort) and with low costs (thanks to cheaper materials and easier fabrication processes), as discussed so far. Aimed at developing such a type of application, the MRI compatibility of the silicone folded DE actuators considered in this work was investigated very recently [22]. For this purpose, an actuator sample was tested within a 3 Tesla scanner (Figure 24.13) operated

Figure 24.13 Testing of the MRI compatibility of a silicone folded actuator arranged, in combination with a standard saline phantom, inside a three-Tesla scanner.

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with gradient-echo echo-planar imaging (GE–EPI) sequences. Specifically, two types of fundamental tests were performed. Firstly, the stress-field characteristic of the actuator was measured both inside and outside (control condition) the MRI scanner operated with GE–EPI sequences; in each case, both a static (step-wise) and a dynamic (sinusoidal) driving of the actuator was tested [22]. Secondly, a signal-to-noise ratio (SNR) investigation was performed on images acquired from a standard phantom (saline solution) during activations of the actuator both inside and outside (control condition) the MRI scanner operated with GE–EPI sequences; in each case, both a static (step-wise) and a dynamic (sinusoidal) driving of the actuator was tested [22]. The static active stress measured outside and inside the MRI scanner is plotted as a function of the applied electric field in Figure 24.14a. The evident lack of any significant difference between the data series suggests a negligible effect of the applied magnetic field and the GE–EPI sequences on the actuator response. Similarly, the performance of the actuator is not affected by the MRI environment and imaging sequences during a sinusoidal driving, as shown by the example reported in Figure 24.14b.

(a)

1.6 Outside MRI scanner Inside MRI scanner (EPI Imaging)

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(b)

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0

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Figure 24.14 Comparison of the response of the silicone folded actuator inside and outside the MRI scanner operated with GE–EPI imaging sequences: (a) static driving; (b) sinusoidal driving.

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Figure 24.15 Comparison of phantom images acquired during static activations of the silicone-made folded actuator: (a) control image recorded with the actuator outside the scanner; (b–e) images obtained while driving the actuator inside the scanner at 0, 4, 6 and 8 V/mm, respectively; (b’–d’) subtraction of the control image (a) from the images (b–e).

MRI images acquired from the phantom during activations of the actuator are presented in Figure 24.15. SNR analyses performed on these images reported no significant losses resulting from the introduction and operation of the folded DE actuator inside the MRI scanner. Such a result was obtained for both static and driving tests [22]. The combination of these results provides a demonstration of the MRI compatibility of the actuators considered. Accordingly, this outcome encourages the development of MRI-compatible, easily wearable and cheap hand splints equipped with such a type of polymer device. 24.9.2

Electromyography-Controlled Hand Splint

Another route of envisaged development of the considered hand splint is represented by its integration with electromyography (EMG). The latter consists of recording bioelectric signals generated by neuromuscular activity. As such, EMG signals are an electrical display of neuromuscular activations associated with contractions of skeletal muscles, regulated by the nervous system. In general, providing a motor rehabilitation system with a so-called myoelectric controller, that is a control unit fed with EMG signals, has the potential to offer multiple opportunities for improving the functionality and the versatility of the rehabilitation system and, thus, the clinical efficacy of the rehabilitation procedures that can be accomplished with it. EMG potentials are today largely used as control inputs for myoelectrically based powered systems designed for different body parts. Focusing our attention on the hand, the literature offers a large number of studies concerned with EMG controlled hand orthoses (see, for instance, [23–25]), in addition to the equally relevant (although not of interest in this chapter) hand prostheses (see, for instance, [26–28]). EMG measurements can be performed noninvasively and comfortably for the user by means of surface electrodes (also known as skin or cutaneous electrodes). This feature makes EMG technology easily and readily applicable for controlling robotic devices [29]. Such a use is also favoured by the relative ease of detecting EMG signals, due to their quite high amplitudes that can reach the order of 1–10 mV.

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Figure 24.16 Schematic drawing of the concept of a myoelectrically controlled hand splint. EMG signals are recorded by means of surface electrodes and are processed by a myoelectric controller that drives the actuators of the splint.

To control orthotic (and also prosthetic) systems for the hand, EMG biopotentials can be advantageously captured, if possible, from the forearm. Figure 24.16 depicts such a case. The most elementary methodology for EMG based controls relies on the extraction of the signal’s amplitude or rate of change from one single channel. More information can be obtained by using two channels, corresponding to two primary electrodes placed on two antagonist muscles. The neuro-muscular signals generated by the user have to be detected, elaborated and used by the machine, approximately in real-time, to perform specific tasks. As an example, the control unit can either facilitate or hamper the execution of a certain rehabilitation exercise, according to the actual neuro-muscular inputs generated by the user. For this purpose, the EMG inputs can trigger a different response of the controller, which can react by increasing or decreasing the voltage supplied to the actuators. As a result, the rehabilitation exercise can comply with the patient’s dynamically varying needs, continuously defined by the controller, according to both its control strategy (algorithm) and the actual EMG evolution. This evidently offers different opportunities with respect to exercises based on pre-programmed modalities. Aimed at driving the device, the controller has to process the EMG signal, typically according to the following successive actions: (i) action onset detection (i.e. identification of the instant when the muscle goes from the relaxed to the contracted state); (ii) feature extraction; (iii) pattern classification. Several algorithms can be used for detecting the movement onset [30, 31] and for extracting features and classifying related patterns [28, 32]. Real-time pattern discrimination and classification is certainly one of the most delicate issues for EMG signals (as for any other type of bioelectric information); with respect to this, neural network based algorithms are today largely adopted [26, 28]. In order to develop EMG based control systems for DE actuators, preliminary investigations have been performed [33]. As an example of the outcomes of such initial studies, Figure 24.17 reports signals related to an EMG based activation of a folded contractile actuator. An example of a myoelectrically triggered contraction of the device is shown in Figure 24.18.

Dynamic Splint-Like Hand Orthosis for Finger Rehabilitation 2.0 Pre-processed EMG [mV]

Recorded EMG [mV]

3 2 1 0 –1 –2 –3

459

500

1000 1500 2000 Time [n° of samples] (a)

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1000 1500 2000 Time [n° of samples] (b)

2500

0

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–1 –2 –3 –4 –5 –6 –7

500

1000 1500 2000 Time [n° of samples] (c)

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Figure 24.17 Example of real-time activation of a silicone folded actuator driven by electromyography: (a) raw EMG signal as detected; (b) elaboration of the EMG signal with a levelling above a certain threshold, used as an input to drive the actuator; (c) contraction of the folded DE actuator in response to an amplified replica of the signal in (b).

Figure 24.18 Example of an EMG driven contraction of a silicone folded actuator.

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24.10 Conclusions This chapter has presented ongoing research activities aimed at developing dynamic hand splints with electrically modulated compliance. Preliminary results suggest the feasibility and efficacy of the proposed concept based on the use of contractile dielectric elastomer actuators. Rehabilitation orthoses equipped with such actuators could offer several potential advantages over alternatives based on conventional actuation technologies; the most significant benefits include lightness, flexibility, comfort, wearability, portability and lack of noise, along with low cost. Therefore, orthotic systems endowed with dielectric elastomer actuation have the potential to open new paradigms in the field of wearable mechatronic systems for rehabilitation. Future developments should be aimed at developing actuators with improved performances, in order to enlarge the admissible working range of the hand splint. Moreover, an MRI-compatible version of the system and an EMG controlled one are envisaged as further parallel developments.

References 1. Phoenix outrigger. Available online at: www.splinting.com. Website visited on 15 October 2008. 2. LMB Wrist Extension Assist. Available online at: www.homecraft-rolyan.com. Website visited on 15 October 2008. 3. Rolyan adjustable outrigger. Available online at: www.homecraft-rolyan.com. Website visited on 15 October 2008. 4. DeROM Dynamic Range of Motion Wrist Splint. Available online at: www.sammonspreston.com. Website visited on 15 October 2008. 5. Tondu, B. and Lopez, P. (1997) The McKibben muscle and its use in actuating robot-arms showing similarities with human arm behaviour, Industrial Robot, 24, 432–9. 6. Sasaki, D., Noritsugu, T., Takaiwa, M. and Yamamoto, H. (2004) Wearable Power Assist Device for Hand Grasping Using Pneumatic Artificial Rubber Muscle, Proceedings of the 2004 IEEE International Workshop on Robot and Human Interactive Communication, Kurashiki, Okayama, Japan, 20–22 September 2004, 655–660. 7. Bouzit, M., Burdea, G., Popescu, G. and Boian, R. (2002) The Rutgers Master II – New Design Force-Feedback Glove, IEEE/ASME Trans. Mechatronics, 7, 256–63. 8. Benjuya, N. and Kenney, S.B., (1990) Hybrid Arm Orthosis, J. Prosthetics and Orthotics, 2, 155–63. 9. Makaran, J., Dittmer, D. K., Buchal, R. O. and MacArthur, D. (1993) The SMART wrist hand orthosis (WHO) for quadriplegic patients, J. Prosthetics and Orthotics, 5, 73–6. 10. Khanicheh, A., Muto, A., Triantafyllou, C., et al. (2006) fMRI-compatible rehabilitation hand device, J. Neuroeng. Rehab., 3, 3–24. 11. Pelrine, R. E., Kornbluh, R. D. and Joseph, J. P. (1998) Electrostriction of polymer dielectrics with compliant electrodes as a means of actuation, Sensors and Actuators A: Phys., 64, 77–85. 12. Pelrine, R. E., Kornbluh, R. D., Pei, Q. and Joseph, J. P. (2000) High-speed electrically actuated elastomers with strain greater than 100%, Science, 287, 836–9. 13. Carpi, F., De Rossi D., Kornbluh, R., et al. (Eds), Dielectric elastomers as electromechanical transducers. Fundamentals, materials, devices, models & applications of an emerging electroactive polymer technology, Elsevier Press, Amsterdam, The Netherlands, 2008. 14. Carpi, F., Salaris, C. and De Rossi, D. (2007) Folded dielectric elastomer actuators, Smart Mat. Struct., 16, S300–5.

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15. Braddom, R.L. Physical Medicine and Rehabilitation, WB Saunders, Oxford, 2006. 16. Salter, M. and Cheshire, L. Hand Therapy: Principles and Practice, Butterworth-Heinemann, Oxford, 2000. 17. EMCO high voltage corporation. Available online at: www.emcohighvoltage.com. Website visited on 15 October 2008. 18. Carpi, F., Gallone, G., Galantini, F. and De Rossi, D. Enhancing the dielectric permittivity of elastomers, in Dielectric elastomers as electromechanical transducers. Fundamentals, materials, devices, models & applications of an emerging electroactive polymer technology (Carpieds, F., De Rossi, D., Kornbluh, R. et al.), Elsevier Press, Amsterdam, The Netherlands, 2008. 19. Carpi, F. and De Rossi D. Contractile monolithic linear actuators, in Dielectric elastomers as electromechanical transducers. Fundamentals, materials, devices, models & applications of an emerging electroactive polymer technology, (Carpieds, F., De Rossi, D., Kornbluh, R. et al.), Elsevier Press, Amsterdam, The Netherlands, 2008. 20. Carpi, F., Gallone, G., Galantini, F. and De Rossi, D. (2008) Silicone-Poly(hexylthiophene) blends as elastomers with enhanced electromechanical transduction properties, Adv. Funct. Mat., 18, 235–41. 21. Tsekos, N. V., Khanicheh, A., Christoforou, E. and Mavroidis, C. (2007) Magnetic resonance– compatible robotic and mechatronics systems for image-guided interventions and rehabilitation: a review study, Ann. Rev. Biomed. Eng., 9, 14.1–14.37. 22. Carpi, F., Khanicheh, A., Mavroidis, C. and De Rossi, D. (2008) MRI Compatibility of Silicone Made Contractile Dielectric Elastomer Actuators, IEEE/ASME Trans. Mechatronics, 13, 370–4. 23. Mulas, M., Folgheraiter, M. and Gini, G. (2005) An EMG-controlled Exoskeleton for Hand Rehabilitation Proceedings of the 9th IEEE International Conference on Rehabilitation Robotics, Chicago, IL, 28 June–1 July 2005, 371–4 . 24. DiCicco, M., Lucas, L. and Matsuokd, Y. (2004) Comparison of Control Strategies for an EMG Controlled Orthotic Exoskeleton for the Hand Proceedings of the 2004 IEEE International Conference on Robotics and Automation, New Orleans, LA, 26 April–1 May 2004, 1622–7. 25. Lucas, L., DiCicco, M. and Matsuoka, Y. (2004) An EMG-controlled hand exoskeleton for natural pinching, J. Robot Mechatronics, 16, 482–8. 26. Castellini, C. and Van der Smagt, P. (2008) Surface EMG in Advanced Hand Prosthetics, Biological Cybernetics, 18 November 2008 (e-publication). 27. Bitzer, S. and Van der Smagt P. (2006) Learning EMG control of a robotic hand: Towards active prostheses, Proceedings of ICRA, International Conference on Robotics and Automation, Orlando, FL, 15–19 May 2006, 2819–23. 28. Zecca, M., Micera, S., Carrozza, M. C. and Dario, P. (2002) Control of multifunctional prosthetic hands by processing the electromyographic signal, Critical Rev. Biomed. Eng., 30, 459–85. 29. Navarro, X., Krueger, T. B., Lago, N., et al. (2005) Critical review of interfaces with the peripheral nervous system for the control of neuroprostheses and hybrid bionic systems, J. Peripheral Nervous System, 10, 229–58. 30. Micera, S., Vannozzi, G., Sabatini, A.M. and Dario, P. (2001) Improving detection of muscle activation intervals, IEEE Eng. Med. Bio., 20, 38–46. 31. Micera, S., Sabatini, A. M. and Dario, P. (1998) An algorithm for detecting the onset of muscle contraction by EMG signal processing, Med. Eng. Phys, 20, 211–5. 32. Crawford, B., Miller, K., Shenoy, P. and Rao, R. (2005) Real-Time Classification of electromyographic signals for robotic control, Technical Report No. 2005-03-05, Department of Computer Science, University of Washington. 33. Carpi, F., Raspopovic, S. and De Rossi, D. Activation of dielectric elastomer actuators by means of human electrophysiological signals, in Smart Structures and Materials 2006: Electroactive Polymer Actuators and Devices (ed. Y. Bar-Cohen), Proceedings of SPIE, 6168, 61681B-1–7, 2006.

Index Note Pl denotes color plate Acrylic elastomers 389–91 performance and characteristics 390–1 thickness strain response 327 transducer configurations 391–2 transverse strain response 327 vs silicone rubber 389–91 AC strain response term (Delta x) 323 Actin–myosin muscle system 7, 26 Agarose 11–12 Albumin exclusion 51 Texas Red labelled 51 Alpha-cyclodextrin 49 Aminobenzophenone 48 Aminoethylmethacrylate 21 Ammonium and potassium counter-ions 21 Aneurysm surgery (in brain), microcatheters 124 Animal structures, marine gels 8–9, 26, 27 Ankle prosthetic device 408 Antigen–antibody binding 46 Antigen-responsive membrane, gating properties 46 Antimicrobial system, self assembly triggered inside microbe with own enzymes 54 Aortic blood flow, peristaltic compression 156 Arterial thrombotic events, see Atherosclerosis Arteries, catheters for thrombosis sample exfoliation 357–68 Artificial muscles and sensors 138, 141, 149–59, 196 actuation and sensing capabilities 147 endoscopically surgically insertable configuration for IPMNC polymeric muscles 157 epoxy hydrogels 17

implantable functional replacements for natural muscle (EPAMs) 403–5 ankle joint 407–8 arm 406–8 artifical diaphragm 403–5 cardiac muscles 138, 149–59 facial muscles 405 peristaltic compression on aortic blood flow 156 limb prosthetics and orthotics 406–8, 443–63 target performance characteristics 26–7 transducers, universal muscle actuator (UMA) 396 see also Heart compression device Atherosclerosis drug-eluting stent (DES) implantation 339 economic burden, US 337 IVUS 336–56 positive remodeling attenuates encroachment of plaque into lumen 338 vulnerability of a lesion to rupture 338–9 Atom-transfer radical polymerisation (ATRP) 76 Azobenzene, photoisomerization 23 Barium ferrite particles, carrageenan hydrogels 25 B-cyclodextrins (b-CDs) 75 Beta-galactosidase, COS-1 cells, dNA complexed with P(NIPAM)-co-DMAEMAcobutylmethacrylate 78 Bilaminar membrane hydrophones 374 Bilayer actuators 254 cell-based sensors 256–7 cell manipulation 255–6 drug delivery 254–5 single hinge, flaps and lids 253–4

Biomedical Applications of Electroactive Polymer Actuators Edited by Federico Carpi and Elisabeth Smela © 2009 John Wiley & Sons Ltd. ISBN: 978-0-470-77305-5

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Bimorph beam model IPMC actuators 181 MSC/ NASTRAN 182 Biocytin (substitute for biotin) 48 Bioelectronic nose 257 Biological gels 8–10 Biological odorant sensors 256–7 Biopsies, prostate cancer 412–14 Biorecognition motifs 54, 56 enzyme cleavable 54 peptide sequences 46 Biotin, as linker between PPy surface and therapeutic 80 Biped walking robots 134–5 Bis(trifluoromethanesulfonyl)amide (TFSA) 207 Bis(trifluoromethanesulfonyl)imide (TFSI) 205 Black box models 113 Blood pressure, pulse rate and rhythm measurement 144 Blood vessels catheters for thrombosis sample exfoliation 357–68 simulators 124–5 Bovine serum albumin (BSA) protein, aTRP 77 Brachytherapy, prostate cancer 412–14 Bragg’s law 32 Braille displays 127–30, 265–77, 427–43 actuator arrangement suitable for Braille pin 268 calculated actuation length changes 270 design and fabrication technique 431–7 essential and desirable characteristics of electronic Braille cells 268 IPMC actuators 275 polypyrrole actuators for Braille pins 271–4 requirements for electronic Braille screen 266–7 small-scale vibratory motion or linear displacement 428 virtual reality tactile display (VRTD) 440–1 wearable tactile display device 164–6, 437–9 Brain, aneurysm surgery 124–5 Broadband transmitters, pVDF, calibrating hydrophones 374 Bucky gel electrodes 117 1-butyl-3-methylimidazolium (BMI) 207 Calcium alginate 11 Calcium ion concentration, drive changes in molecular conformation 22 Calcium-responsive hydrogels 46 Calmodulin (CaM) 51 calcium-bound VS ligand-bound 53 Cancer therapy, polymer—drug conjugates 67

Cantilever, mechanical deflection, UV irradiation 23 Capacitors, transduction 428 Carbon nanotubes, singlewalled (SWNTs) 117 Cardiac syndromes 337–56 see also Atherosclerosis Cardiac therapy 137–60 actuation and sensing capabilities of IPMNCs 147 see also Heart cells; IPMNCs Cardiomyocytes dielectric elastomer actuators 408–9 scaffold degradation 66 Carrageenan hydrogels, barium ferrite particles 25 Cartilage articular 9 artificial cartilage 12 Catheters history and applications 229–31 thrombosis sample exfoliation 357–68 see also Microcatheters Catheterscope, active tip bending system 236 Cell culture assays, dielectric elastomer actuators 408–9 Cell manipulation 255–6 cell-based sensors 256–7 Cell-responsive hydrogels 46 Cell/tissue/organ based therapies, tissue engineering and regenerative medicine strategies 65–6 Ceramics thickness strain response 327 transverse strain response 327 Charge coupled device (CCD) camera 362 Chemically driven gel actuators 20–2 Chiral polymers 358 Chymotrypsin-responsive hydrogels 46, 54 Cluster networks 142 Cochlear implants 244–5 Collagen belt, contraction and expansion 7 Colloidal crystals 32 Complementary metal oxide semiconductor (CMOS) process 341 circuitry 257, 260 Composite gels 12–13 Conconcanavalin A (Con A) 46 Conducting polymers, see Conjugated polymers (CPs) Conjugated polymers (CPs) 195–227, 246 bioerodable and biodegradable 284 biomedical applications 282–3 carriers for drug molecules 285 charge transfer processes 208–9 classification 2–3

Index electrochemistry 201–4 applied potential 204 cyclic voltammetry 202–3 electrolyte conductivity 204 ion diffusion 204 polymer conductivity 204 electronic charge injection rate 209 examples 196 infusion micropumps 175 lifetime 260 MEMS sensors 249–64 molecular dopant 197–8, 237, 251 nanostructured 280–5 polymer morphology and geometry 204 properties 195–6, 201, 249–50 vary with choice of dopant and synthesis conditions 199 as smart materials for actuation 1–4 vs IPMCs 246 Contact lens gels 12 Controllable flow rate pump 89–98 Copper chlorophyllin, NIPAM gel 23 Cornea 9 Coronary artery imaging, IVUS 336–56 Coulombic force 323 Cross-linked gels 21 Cyclic loading tests 14 Cyclic voltammetry, conjugated polymers 202 Deep reactive ion etching (DRIE) 344 Deformed shapes of IPMC diaphragms, Pl6 Device design and optimization 218–22 actuator Performance 217–18 Dexamethasone (Dex) 79 implanted poly(ethylvinyl) acetate strips 285 insertion of electrodes 285 PEDOT and PLGA nanotube delivery 285–6 Diabetes, treatment with glucose responsive polymers 49 Diaphragms 181, 182–9 artifical implantable 403–6 for infusion micropumps 183 Dielectric elastomer actuators (DEAs) 387–96 actuation principle 428–9, 447–8 linear contractile elastic actuators 448–9 actuator lengthening compliance–voltage plots, different amplitudes 454 force–voltage plots 453 as basis of tactile display 428 best performing materials 389–91

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binary, for prostate cancer detection and treatment 411–23 biomedical devices 397–410 cell culture assays 408–9 diaphragms 403–5 dynamic hand splints 446–8 facial muscles 405 MRI guided robotic surgery 411–27 pumps 396–9 thickness-mode actuation 400–3 design and fabrication technique 430–2 dimensioning the actuators 449–50 examples 3 initial thickness and the strain in the thickness direction 429 linear contractile (folded actuator) 447 operational considerations 392–3 principles of operation 388, 392–3 safety 391–2 transducers 388–9 configurations 392 Dimethylaminoethyl methacrylate (DMA) 56 Disulfide cross-links 21 DNA condensation coil–globule transition lower critical solution temperature (LCST) 74 rods and toroids 70 polycations 70 DNA–polymer complexes cellular uptake 70–2 transfection efficiency 72 DN gels, stress–strain curves 13 Dopant ions 197–9, 237, 251 choice 204–5 Double network gels 13–14 Drug delivery systems 61–89, 279–99, 301–16 actuated polymers 78–80 barriers to delivery 64–8 bilayer actuators 254–5 controllable flow rate pump 89–101 current delivery methods 285 electropolymerized PPy nanotubes 282 in presence of Dex 78 gels 31–2 implanted vehicles or depots 63 inappropriate or unsuitable drugs 66 intelliDrug system 301 oral route 62–3 oral via buccal mucosa 301–16 parenteral drug delivery 63 therapeutic window in drug therapy 302 topical/transdermal delivery 63 Trojan Horse analogy 68

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Drug-eluting stents (DES) implantation 339 thermo-responsive gels 32 Drying time 15–16 Elastic energy density 326 Elastic modulus 9, 10 Electrically driven appliances, safe current limits 231 Electrically driven gel actuators 8, 23–4 Electroactive chemically cross-linked hydrogels 56–7 Electroactive Polymer Actuators and Devices 2008 Conference (SPIE) 387 Electroactive polymer actuators (ionic EAPs) 195–227, 237, 249–64, 250–3 4 processes leading to volume change and actuation 200 actuated pins for Braille displays 265–77 actuation mechanism 251 actuation strain 250 out-of-plane 259 actuation stress gel of known elastic modulus 27 linear actuators 250–1 applications 30–3, 127–32, 175, 195–6, 235 active steerable microcatheters 237–43 artificial muscles 138, 141, 149–59, 196 biological odorant sensors 257 catheterscope 236 intravascular neurosurgery 127 linear actuators and benders 195–6 microelectromechanical systems (MEMS) sensors 122 micropumps 175–91 microrobots 130–4, 196, 214, 219–21 microscale robot arm 257–8 multi-bilayer devices 257–9 OCT fast scanning catheter 239–40 out-of-plane actuation strain 25 positioning tools 257–9 robotic actuation 142–5 tactile stimulation/displays 161–74 chemical/mechanical processes involved in actuation 198 cochlear implants 244–5 compact tactile display 162 composition, geometry and electrolyte, effect on actuation 204–12 contractile linear actuator attached in series with a restoring spring 213 decreasing actuation occurring with higher isotonic stresses 215 device design and optimization 217–22

doping 204–5, 237, 251 driving active catheters 238 future prospects 222–3 integration into more complex microsystems 260 mechanical system response 212–17 microfabrication 251–3 modelling control of actuation 197 molecular mechanisms of actuation 197–8 operating against a restoring spring 268–9 performance 250–1 various ICPs 200–3 pump systems w. concentric cylindrical layers of ICPs 196 RC time of system 199 redox switching 4, 197 speed of switching between expanded and contracted states 250 see also Bilayer actuators; Conjugated polymers (CPs); Ionic polymer–metal composite (IPMC) actuators Electroactive polymer artificial muscle (EPAM), see Artificial muscles and sensors Electroactive polymers (EAPs) classification 2–3 see also Dielectric elastomer actuators (DEAs); Electroactive polymer actuators (ionic EAPs); electrostrictive polymers; Ionic polymer–metal composite (IPMC) actuators; Piezoelectric polymer film/fibres Electrocutaneous stimulation 162 Electrodes bucky gel 117 diaphragms for infusion micropumps 183 flexible wire 28 gold pseudo-reference electrode 259 insertion cochlear implants 244–5 dexamethasone 285 patterned 124 plating ion exchange membranes 105–6 gold 106, 108 platinum 104 polyimide 285 puffer probes, microfluidic channels inside electrode shank 285 Electromagnetic DC motors 446 Electro- and magneto–rheological composites 25 Electromechanical (E-M) materials 319–30 elastic responses 326 electromechanical coupling factor (k) 325 frequency and temperature dependence 326 properties 324–5

Index Electromyography (EMG) biopotentials, capture from forearm 458 EMG-controlled hand splint 457–9 EMG-driven contraction, silicone folded actuator 459 Electrophoresis, SDS–PAGE 77 Electropolymerisation 79, 282 Electrospinning 286–8 Electrostress diffusion coupling (ESDC) model 114 Electrostrictive effect 319–34 electromechanical coupling factor (k) 325 Electrostrictive polymers 330–2, 388–9 examples 3 E-M polymers, see Conjugated polymers (CPs); Ionic polymer–metal composite (IPMC) actuators Endocytosis, schematic 71 Endolysosomes 71–2 Enhanced permeation and retention (EPR) effect 67, 75 Enzyme-cleavable linker (ECL) 50 Enzymes chemo-, regio- and enantio-selective mechanisms 50 reversible self-assembly of gels 55 Epoxy hydrogels small artificial muscles and sensors 17 swelling–collapse behavior 17 1-ethyl-3-methylimidazolium (EMI) 207 3, 4-ethylenedioxythiophene (EDOT) 288–9 Extracorporeal shock wave therapy (ESWT) 374–7 Eyelid reanimation, artificial muscles and sensors 405 Facial muscles, dielectric elastomer actuators 405 FDA, release of medical products regulations 232 Ferroelectricity 328–30, 357 Ferroelectric polymers poling 358 PVDF–TrFE 357–8 Fibre-reinforced gels 8 Finger rehabilitation 443–62 Fingers, mechanoreceptors 165 Flemion, see Nafion, Flemion (IEPAs) Fluorescence resonance energy transfer (FRET) 50 Focused Ion Beam (FIB) instrument 289 Forisomes 22, 26 Free radical polymerization 16 Froude number 220

467

Gel actuators 7–41 drug delivery applications 78–9 methods 19–26 bi-stable switching 23 chemically driven 20–2, 28 electrically driven 23–4, 27 electro- and Magneto- rheological Composites 25 light Actuated Gels 23 liquid crystalline (LC) elastomers 26 oscillating chemical reactions 22 thermally driven 19–20 performance of muscle 26–30 Gelation, synthetic ionic polymers 18–19 Gel-based machine 15 Gel diffusion coefficient (D) 90 Gel-in-gel thermally driven light modulator 31 Gel microcantilever, in water, before/after UV irradiation 17 Gels chemical and physical formation 16–18 contact printing 18 diffusion coefficients of small solutes 15 dilute, swelling pressure exerted 27 driven by solvent-induced contraction 27 effect of electric field 29 elastic modulus 9, 10 fracture 11 fracture toughness 14 known elastic modulus, actuation stress 27 lightly cross-linked 14 loopers and swimmers 24 oxygen-sensitivity problem 16–17 samples formed by UV irradiation 16 tensile strength 9, 10–11 transport properties 14–15 two-gel multi-layer system 28 vs elastomers 11 Gel sensors 32 Gel swelling, kinetics 19 Gel valves and pumps 30, 97 Gene delivery polymers for 69 design 73 Glucose oxidase (GOx) 49–50 Glucose responsive polymers, treatment of diabetes 49–50 Gold plating, see Electrodes Gradient-echo echo-planar imaging (GE–EPI) sequences 456 Grafted elastomer, thickness strain response 327 Griffith equation 11

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Index

Hall effect isotonic transducer 90 Hand splints 443–62 active dynamic hand splints 445–6 classification 444 dynamic splint with dielectric elastomer actuators (DEAs) 446–8 future developments EMG-controlled hand splint 457–9 MRI-compatible hand splint 454–7 linear contractile elastic actuators 448 passive dynamic hand splints 444–5 prototype 450–4 splint mechanics 449–51 Heart cells dielectric elastomer actuators 408–9 scaffold degradation 66 stretching, using planar EPAM actuators 409 Heart compression device 149–59 bradycardic (pacing) and tachyarrhythmic (cardioverting/defibrillating) control 151 compression band, Pl2, 153, 154 equipped with IPMNC fingers 151 mini heart compression device 153–4 multi-fingered heart compression device (MFHCD) 155–7 preparation of thick IPMNC fingers 155–7 pressure generation vs electrode thickness, Pl3, Pl4 154–5 supported on slender flexible stem 152 synchronism with LV contraction 149 transcutaneous recharging of implanted batteries 149–50 see also Artificial muscles and sensors Heart failure (congestive, CHF) 137, 139–56 Heart muscle, properties 138 Heart valves 139 HEMA–gelatin gels 14 Hexafluorophosphate (PF6) 207–8 High frequency ultrasound 335, 340–56 High intensity focused ultrasound (HIFU) surgery 374, 377 Hooke’s law 320 Hydrofoil, variable camber 196 Hydrogelation control 53 Hydrogel matrix, kinetic equation of motion 90 Hydrogels, bio-responsive 43–57 actuation 90–1 changing cross-linking density 46–9 changing electrostatic repulsion/ attraction of polymer chains 49–50 conformational changes 51–3 antigen-responsive 48

applications 31 bending curvature 29 calcium-responsive 46 cell-responsive 46 chemical cross-linking 44, 46–7 chymotrypsin-responsive 46 cross-linking vs Electrostatic examples 47 defining 44–5 electroactive chemically cross-linked 56–7 epoxy, small artificial muscles and sensors 17 glucose responsive, and diabetes 49–50 as osmotic pumps 89, 97 physical 44, 53–6 ‘smart’ 44, 66, 73–8, 89–101 Hydrogen-bonding polymers 12 Hydrophones, piezoelectric 369–77 applications diagnostic ultrasound 372–3 extracorporeal shock wave therapy (ESWT) 374 therapeutic ultrasound 374–6 bilaminar membrane 374 spot-poled membrane 371 spot poling 371 Hydroxyethylmethacrylate 12, 21 Hydroxypropylcellulose (HPC) 20 Imaging, high frequency ultrasound 335–56 Implanted vehicles or depots 63 Impregnation reduction (IR) 106–7 Infusion micropumps 175–91 characteristics as actuator and sensor 177 control system (Renesas M16C microcomputer) 180 diaphragms 181–9 circle-shaped diaphragm vs square-shaped diaphragm 186 circle-shaped IPMC diaphragm 188 normal mode analysis 186–8 pressure effect on stroke volume 188 schematic 189 shapes of electrodes 183 interdigitated 20-finger IPMC 180 modelling 181 performance 181 prototype IPMC pump 178 pump chamber cover 178 schematic 189 IntelliDrug system 301–16 assembly and clinical testing 314–15 exploded illustration 305 osmotic pressure pump 305 schema 304 valve concepts 310–14

Index Intelligent Polymer Research Institute 244 Interchain bonding, self-assembly of gels 18 Intravascular neurosurgery 127 Intravascular ultrasound (IVUS) 335–56 commercially available 340 focusing methods 340–1 lateral resolution, axial resolution and depth of field 340 polymer transducers 341–4 testing 345–7 rotating IVUS geometry 336 automated quantitative analysis 336–7 Ion exchange membrane, electrode plating 105–6 Ionic conductive polymer gel film (ICPF) 105 Ionic liquids, solvent-less electrolyte 207 Ionic polymer conductor composites (IPCCs) 140 Ionic polymer conductor nanocomposites (IPCNCs) 140 Ionic polymer–metal nanocomposites (IPMNCs) 137–57 Ionic polymer–metal composite (IPMC) actuators 29, 103–19 applications 116–17, 122, 127–32 cardiac muscle therapy 137–60 catheterscope 236 intravascular neurosurgery 127 microcatheters 124–6, 234–46 microelectromechanical systems (MEMS) sensors 122 micropumps 175–91 microrobots 130–4, 196, 214, 219–21 robotic actuation 142–5 tactile stimulation/displays 161–74 background 176–7 basic movements 122 bimorph beam model 181–2 diaphragms 181–9 dry actuator of bucky gel 117 fabrication 104–8, 122–4, 235–6 impregnation reduction (IR) 106–7 microactive guide wire 235 reductant permeation (RP) 106 importance of the Young’s modulus 213–14, 217, 254 IPMC sensor/actuator system 124 Measurement 108–10 characteristic time of back motion 116 models 113–16 black box model 113–14 electrostress diffusion coupling (ESDC) model 114 nanocomposites (IPMNCs) 137–60 3-Dimensional fabrication 141–2 actuation and sensing capabilities 147

469

artificial muscles 141, 149–59 displacement characteristics of an IPMNC 143 electrically-induced robotic actuation 142–5 four-fingered compression system 142 frequency dependent dynamic deformation characteristics 143–4 generation of power under normal pressure 144–5 long cycle oscillation vs blocking force 147 material testing system 145 modeling and simulation 146–8 movement capabilities suitable for heart compression 149 oscillatory mechanical excitations 145 output voltage due to normal impact of a 200 N load 146 peristaltic compression on aortic blood flow 156 properties 141 sensing and transduction properties 144 theory of actuation mechanisms 147 tip blocking force and the associated deflection 144 see also Ionic polymer–metal composite (IPMC) actuators normal mode analysis, iPMC diaphragm, Pl7 performance properties 110–13, 176 predicting behaviour of IPMC diaphragms 181–2 properties, strain rate and stress 235 restoring spring stiffness 213 vs conjugated polymers (CPs) 246 see also Conjugated polymers (CPs); Ionic polymer—metal nanocomposites (IPMNCs) Ionic selfassembly, sequential dipping into anionic and cationic polymers 18 Ion transport, mechanoelectric effect 147 IPMNCs 140–1, 144, 145, 146, 147, 149–55, 156 IVUS transducer, coronary artery imaging 351 J-shaped tensile stress–strain curve Keto to enol tautomerization Kroneker delta 91

9, 14

23

Landau model, phase transitions 23 Laser cutting machine; CAD design for interdigitated IPMC, Pl5 Laser displacement meter 109

470

Index

Left-ventricular assist systems (LVAS) 137–60 examples 138 Legumes, specialized cells control of fluid flow 26 Lenses, fabrication of miniature lenses 340–1 Leucocyanides 23 Light actuated gels 23 Light modulator, gel-in-gel thermally driven 31 Limb prosthetics and orthotics 406–8, 443–63 Linear actuators and benders 195–6, 250–1 Liposomes, polycation poly-L-lysine 69 Liquid crystalline (LC) elastomers 26 Lithotripsy, extracorporeal shock wave therapy (ESWT) 374–6 Lower critical solution temperature (LCST), coil-to-globule transition 74 Lysosomes 71 Magnetic resonance imaging (MRI) systems 411–26 MRI-compatible actuators 455 MRI-compatible hand splint 454–6 MRI guided robotic surgery 411–13 closed-bore MRI manipulation 415 elastically averaged parallel manipulator 412, 415–16, 417–20 precision 416–17 results 420–3 Marine gels 8–9, 26, 27 Master–slave system 173 Material testing system (MTS) 145 Maxwell (stress) effect 324 based polymers 332 Mechanoelectric effect 147 Mechanoreceptors, in human fingers 165 Meissner’s corpuscle 163, 165 Merkel corpuscle 165 Micelles coated with polyethylene glycol, biocompatibility 75 drug delivery applications 74–5 size 75 Microcatheters 234–46 active guide wire catheters 124–6 active steerable catheters 234–46 CP based 237–43 EAP based 235–46 finite element simulation 239 intravascular 238 IPMC based 235–6 non-EAP 234 OCT fast scanning catheter 239–40 rigidity requirement 238–9 SMP based 236 Applications 124–5, 229–31 Biocompatibility 231–2

cost and size 232 design 231–4 passive catheter 233 structural rigidity 232–4 axial (EA), torsional (GJ) and flexural (EI) rigidities 233 Microelectrodes, neural 284 Microelectromechanical systems (MEMS) 249–64 drug delivery 254–6 microfabrication 251–3 sensors 122 test apparatus 345–6 ultrasonic transducer prototype 344 Microfabrication 251–3 Microfibres, rotating fibre driven by external motor 240 Micropumps 175–91, 196 controllable flow rate pump 89–98 Microrobots 196 arm/hand 257–8 underwater 24, 130–4, 196, 219–21 see also Robots Microscale robot arm 257–8 Microtools 257 Microvalve 259–60 Miniature disposable infusion IPMC micropumps, see Infusion micropumps Multi-phase gels 12–13 Multiple bilayer devices 257–9 Muscle as actuator, force 300 kPa 8 Muscle contraction conformation change 8 linear 24 Muscle-like actuators, methods 19–26 Muscles artificial, see Artificial muscles and sensors energy density and power output 26 wasting, artifical diaphragm 403–6 Myocardium, properties 138–9 Myocytes, dielectric elastomer actuators 408–9 Nafion, Flemion (IEPAs) 104–13 applications microcatheters 124–7, 135–46 microrobots 127–36 double-layer capacitance 112 fabrication 122–4 membrane conductivity 107 performance 110–12 response speed and displacement 163 scanning electron micrographs 112 spraying Nafion using an airbrush 177 see also Ionic polymer–metal composite (IPMC) actuators

Index Naltrexone, buccal and intravenous delivery 302–3 Nanofibres electrospinning 286–8 polyaniline (PANi) 282–3 poly(lactide-co-glycolide) (PLGA) 281, 285–90 Nanostructured conducting polymers 279–85 electrospinning 286–8 fabrication methods 280–3 variety of geometries 280 see also Drug delivery systems Nanotubes, PPY and PEDOT 281 Needle hydrophones, piezoelectric ceramic 369–71 Nerve growth factor (NGF), stimulated release 79 Neumann’s principle 320–1 Neural microelectrodes 284 Neurosurgery, intravascular 127 NIPAM gel, copper chlorophyllin 23 NMR, perfluoroionomers 141 N, N’(2-dimethylamino)-ethylmethacrylate (DMAEMA) 77 Nuclear trafficking, complexed DNA 70, 72–3 Nucleic acids, use in therapy 65 Odorant sensors 256–7 Oligodeoxynucleotides, (ODNs) 68 Onsager coefficients 91, 114, 148 Onsager formulation 147 Optical coherence tomography, (OCT) 236 fast scanning catheter 239–40 schematic, actuator set up for OCT 239–40 high speed actuation 242 probe development 236 Optical fibres, scanning 240 Orthotics 406–8, 443–63 Oscillating chemical reactions 22 Out-of-plane actuation strain 259

230,

Pacinian corpuscle 165 Parenteral drug delivery 63 Peltier’s cells 97 Penicillin G amidase 54 PEO–PAA gels 14 PEO–PPO block copolymers 64 Peptide actuators 50 Peptide ampiphiles 53 self-assembly into fibrous cylindrical micelles 54 Peptide bonds, assembling Fmoc-dipeptide methyl ester 55

471

Peptide (de-)phosphorylation 53 Peptides alpha-helix polypeptides 358 and proteins 64–6 short, assembly 46 Peptide sequences, biorecognition motifs 46 Perfluorocarboxylic acid polymer 104–13, 122–5 tubular, gold electroplate, external and internal sides 235–6 Perfluorosulfonic acid polymer 104–13, 122–5, 140 properties 140 water content 106 see also Ionic polymer–metal composite (IPMC) actuators; Nafion, Flemion (IEPAs) Phenylboronic acid, binding of glucose 32 Phosphatase/kinase 54 Phosphated gels 21 Photoresponsive dyes 23 Piezoelectric ceramics bending beams 266 hydrophones 369–77 Piezoelectric constant, PVDF–TrFE copolymer 357 Piezoelectric and electrostrictive effects 319–34 applications 332 elastic responses 326 frequency and temperature dependence 326 materials at room temperature, thickness strain response 327 piezoelectric polymers 328 strain response 322 Piezoelectricity, shear 358 Piezoelectric motion, PLLA fibres 362–3 Piezoelectric polymer film/fibres 358–67 applications 357–8 determining bending motion 361–3 fabrication of IVUS transducers 351 fabrication of prototype system for catheters 363–6 motion 362–7 PLLA fibres 362–7 Piezoelectric PVDF, see Poly(vinylidene fluoride) (PVDF) Plating electrodes, properties 106 Platinum electrodes, plating 103–19 P(NIPAM) materials 74 Polhemus information, magnetic field 165 Poly(2-acrylamido-2-methyl-1-propane sulfonic) acid (PAMPS) 140

472

Index

Poly(3, 4-ethylenedioxythiophene) (PEDOT) 279, 280 controlled release of dexamethasone 291 dispersion polymerization 282 electrochemical deposition of PEDOT nanotubes 288–90 fabrication methods 280–1 properties 284 UV absorption of dexamethasone-loaded PEDOT nanotubes, Pl8 Polyacrylamide 10 Polyacrylamide/poly-N-isopropylacrylamide IPN gels 14 Polyacrylic-acid-bis-acrylamide (PAAM) 140 Polyacrylic acid (PAAC) fibres 7 gels 21, 50 Polyacrylic acid/poly(vinyl alcohol) gels 12, 27 driven electrically 24 strain in bending and weight gain 24 working energy 28 Polyacrylonitrile (PAN) 140 fibres and chemically driven gels 22 Polyaniline (PANi) 196, 237 actuation performance and various properties 201 actuator assembly consisting of 8 PANi twisted fibre yarns inserted in a hollow PANi fibre 200 continuous spun fibres addition of carbon nanotubes 215 as linear actuators 196 drift in actuation strain over time 216 nanofibres rapidly mixed and slowly mixed reactions 283 synthesis 282 Poly(caprolactone) (PCL) 64 Polycation mediated delivery 69 nonviral polycations 70 Polycations, DNA condensation 70 Poly(dimethylsiloxane) (PDMS) 259 silicone rubber 389–91 Polydispersity (PDI) 76 Polyelectrolytes 10, 18 Poly(ethylene glycol) acrylamide (PEGA) 50 Poly(ethylene glycol) diacrylate (PEGDA) 49 Poly(ethylene glycol) (PEG) 67–8 new bioconjugated therapeutics and products in clinical use 68 triblock copolymers 75 Poly(ethylene oxide)–polyacrylic acid double network gels 14

Polyethylenimine (PEI) 69 complexes with DNA 70 entering nucleus 73 release from lysosome 72 Poly(ethylvinyl) acetate, release of dexamethasone from implanted strips 285 Poly(glucosyloxy-ethy1methacrylate) poly (GEMA) 46 Poly(hydroxyethyl methacrylate) (PHEMA) 44 Poly(lactide-co-glycolide) (PLGA) 64, 285–92 controlled release of dexamethasone 291–2 electrospinning process 287–8 nanofibres 285–90 Polylactide (PLA) nanofibre templates 281 Poly-l-lactic acid (PLLA) 358 fibre high shear piezoelectric constant 358–65 motion 362–7 film/fibres, fabrication of prototype system for catheters 363–7 molecular weight (20000) 361 piezoelectric modulus 362 Poly-L-lysine (PLL) 69 Polymer—drug conjugates 67 responsive 76 Polymer–electrolyte systems, redox reactions 201 Polymer gel actuators 7–41 examples 3 properties of gels 8–16 Polymer–metal composites, see Ionic polymer–metal composite (IPMC) actuators Polymer–nucleic acid conjugates 68 Polymer–oligodeoxynucleotides (ODNs) conjugates 68 Polymer–protein conjugates 67–8 Polymers biodegradable natural 66 charge transport, affected by factors 204 ferroelectric 328 poling 358 PVDF–TrFE 357–8 free radical polymerization 16 for gene delivery 69 polymer controlled drug release systems 67–73 salt draining 205 stability against aggregation 63 thickness strain response 327 transverse strain response 327

Index Polymer transducers 341–4 fabrication 342–4 testing 345–7 Polymethacrylic acid (PMAA) gels, changes in pH 21 Poly(N-isopropylacrylamide-co-acrylic acid) P(NIPAM-co-AAc) 48–9, 74, 76–8 gene delivery systems 78 growth, cuBr / bipyridyl system 77 transfection efficiency 77 Poly(N-isopropylacrylamide) (NIPAM) 19, 20 Poly(N-vinylcaprolactam) (PVCL) 20, 27 Polypeptides, see Peptides Polypropylene, non-woven reinforcement 13 Poly(pyrrole) (PPy) 78–80, 196, 237 active element fabrication 241–2 actuation performance and properties 201, 210–12, 307–10 applied isotonic stresses 209 cyclic voltammograms 308 effect of solvent 206–7 exchange of solvent between polymer and electrolyte 207 mechanical output 212–13 PPy(TFSI) films 205–7 rate of actuation 243 relevant information 307 scanning an optical fibre 240 selection of counter electrode 309 biotin as a linker between PPy surface and therapeutic 80 Braille pins 271–4 charge-switchable drug delivery system 79 composition, geometry and electrolyte 204–17 charge transfer processes 208–12 dopant ion used 204–5 platinum wire encapsulated within device during growth 209 porosity/morphology 212–13 and Dex 78 electroactivity PPy(DBS) steady decline 260 redox cycle 237 electrochemical reaction 197 electrolyte conductivities raised 206 resistance compensation techniques 209 electropolymerisation nanotubes 282 in presence of Dex 78 fabrication, nanotubes 281 helix tube actuator 210 lifetime 260 microvalve 259–60 nanotubes 282

473

oral valve 304–9 patterning of polypyrrole electrodes 242 PPy(DBS) in aqueous Na(DBS) electrolyte 210 PPy(DBS)/Au bilayer 254, 259 PPy(DBS) electroactivity, steady decline 260 rate of actuation 243 redox state and the PPy film volume 211 structure 197 Polyrotaxanes 76 Polytetrafluoroethylene (PTFE) 141 Polythiophene (PTh) 196 actuation performance and properties 201 see also Poly(3, 4-ethylenedioxythiophene) (PEDOT) Polyurethane elastomers 11 steerable catheters 236 Poly(vinyl alcohol) gels 12, 27 Poly(vinylidene fluoride) (PVDF) 50, 275, 321, 369–84 applications 369 biomedical ultrasound devices 369–84 broadband transmitters, calibrating hydrophones 374 cavitation sensors 377 receivers 376 sensors, measured pulse fidelity 374 silicon integrated circuit (IC) fabrication 341 membrane 245 permanent dipole moment 328 and ferroelectricity 329 properties 369 thickness resonance 375 PVDF–TrFE copolymer 330–2, 341–4 fabrication to MEMS 341–4 ferroelectricity 328–30, 357 piezoelectric constant 357 thickness strain response 327 transverse strain response 327 TrFE and TFE 328 Poly(vinylimidazole) derivatives 69 Poly(vinyl-methyl-ether) (PVME) 89 fluid–matrix friction coefficient 96 gel dynamics 91–2 pump functioning 97, 98 PVME actuator length calculation 93–4 PVME–sodium alginate emulsion 90 response time constant as a function of temperature 93 shear diffusion coefficient 94 shear elastic modulus 96 stress as a function of temperature 97 stress generated temperature change 95 Young’s elastic modulus (E) 95, 213–14

474

Index

Poly(vinyl sulfonate) (PVS) 205 Potentiostat 109 Printed polyelectrolyte complex filaments 18 Probe development, OCT 236 Propylene carbonate (PC) PC electrolyte associated with parasitic redox reactions 208 solutions 207–8 Prostate cancer 411–24 biopsies and brachytherapies 412–14 MRI guided robotic surgery 411–13 closed-bore MRI manipulation 415 elastically averaged parallel manipulator 415–16, 417–20 precision 416–17 results 420–3 Prosthetics and orthotics 406–8, 443–63 Proteins beta-sheet segments (KLD) 54 conformational changes 51–3 Proteoglycan gels 9, 26 Proton sponge mechanism 72 P-toluene sulfonate 205 Pump, see Micropumps Pyrrole monomer, polymerization 241 Radiation patterns 349 Redox reactions conjugated polymers 203–4, 237 parasitic 208 polymer/electrolyte systems 201 Redox switching 196–7 Reductant permeation (RP) 106 Responsive polymer—drug conjugates 76 Responsive polymer phase transitions, binding interactions with DNA 77–8 Responsive polymer—protein conjugates 76 Reversible self-assembly by use of enzymes, Pl1 RGD, cell binding epitope 56 Rigid-rod polymers 13 Robotic actuation, electrically-induced 142–5 Robotic fish propulsor fin 196 Robots biped walking 134–5 underwater microrobot 130–4, 196, 214, 219–21 see also Microrobots Rubbers crystallization 11 fracture 11 Ruffini endings 165

Salt draining 205 Scaffolds degradation 66 synthetic support structures 66 Scallop, performance of muscle 26–30 Self-assembly of gels interchain bonding 18 reversible by use of enzymes 55 Semi-interpenetrating network (semi-IPN) hydrogel 46 Shaped memory alloys (SMA) 229 Shaped memory polymers (SMPs) 235, 236 steerable catheters 236 Sheet-type Braille display 128–30 Silicone thickness strain response 327 transverse strain response 327 Silicone folded actuator EMG-driven contraction 459 gradient-echo echo-planar imaging (GE—EPI) sequences 456 real-time activation 459 static activation, phantom images 459 testing of MRI compatibility 455 Silicone rubber (polydimethyl siloxane) 389 performance and characteristics 390 vs acrylic elastomers 389–91 Silicon microcantilevers 32 Silicon micromachined phantom 349–50 Single hinge bilayer devices, flaps and lids 253–4 Single-phase gels gel concentration 15 mechanical properties 9–10 Sinoatrial (SA) node 139 Size exclusion chromatography (SEC) 77 Skin, tactile receptors 161–4 ‘Smart’ bio-responsive polymers 1–4, 44, 66, 73–8, 89–101 active steerable catheters 234–46 diblock, forming micelles 75 micellar states 75 microvalve 259–60 performance properties 176 soluble polymers 73–6 synthetic gene delivery vectors 69–78 thermally driven actuator for controllable flow rate pump 89–101 Soft actuators, electrical response 29 Solid polymer electrolyte (SPE) 238 Solvent transport, mechanoelectric effect 147 Spot poling, piezoelectric hydrophones 371–2 Stents, thrombotic events, drug-eluting stent (DES) implantation 339–40

Index Stimulation devices, thickness-mode actuation 400–3 Stokes–Einstein law 114 Streptavidin 76 biotinylated NGF 80 polymeriSation via biotin initiator 77 Styrene/divinylbenzene families 142 Subtilisin 54 Sulfonated gels 21 Swelling film devices 258–9 Swimming device design 219–21 polymer actuators 221 Synthetic ionic polymers, gelation 18–19 Synthetic support structures, scaffolds 66 Tactile display methods to produce virtual touch 161–74, 427–43 advanced applications 437–40 proposed tactile cell 437–40 advantages 163 design and fabrication technique 430–1 dielectric elastomer actuators (DEAs) 428 friction sensation 171–3 pin matrix density 428 pressure sensation 168–9 roughness sensation 169–71 selective stimulation, tactile synthesis 165–7 small-scale vibratory motion or linear displacement 428 synthesis of total textural Feeling 172–3 texture synthesis 167–8 wearable tactile display 164–6, 437–9 Tactile receptors in human skin 161 Taylor series expansion 431 Templated electrochemical growth around nanoscale electrospun filaments 292 polylactide 281 Tensile strength, gGriffith equation 11 Tetrabutylammonium-PF6 207–8 Tetraethyl ammonium (tEA) ion 110 Tetrafluoroethylene (TFE) 328 Thermally driven gel actuators 19–20 for controllable flow rate pump 89–101 Thermally driven light modulator, gel-in-gel 31 Thermolysin 54 Thermomechanical gels, dynamics 91 Thermoresponsive polymers drug eluting stents 32 ppP(MEO2MA-co-OEGMA) and P(NIPAM) 74

475

Thickness-mode actuation 400–3 balance sensors, elderly/diabetic patients 402 massaging applications 401 other applications 403 Thickness strain response 327 initial thickness and strain in thickness direction 429 Thrombotic events catheters for sample exfoliation 357–68 stents 339–40 Timoshenko equation 254 Transducers, configurations 391–2 Transduction mechanical of a two-plate capacitor 428 two-plate capacitor 428 Transglutamase 46 Transverse strain response 327 Trifluoperazine ligand (TFP) 53 Trifluoroethylene (TrFE) 328 Tubular IPMC actuator control of active microcatheter 126 with four electrodes 125 for intravascular neurosurgery 127 Tumour cells, enhanced permeation and retention (EPR) effect 67 Ultrasonic transducers 335–6, 340–56 continuous wave axial radiation patterns 346–8 lateral resolution, axial resolution and depth of field 340 test apparatus and results 345–51 Ultrasound high intensity focused ultrasound (HIFU) surgery 374 hydrophones, piezoelectric diagnostic ultrasound 372–3 therapeutic ultrasound 374–6 Underwater microrobots 130–4, 196, 214, 219–21 Universal muscle actuator (UMA) transducers 396 Urokinase 54 US FDA, release of medical products regulations 232 UV absorption of dexamethasone-loaded PEDOT nanotubes, Pl8 UV irradiation gel formation 16 keto to enol tautomerization, mechanical deflection of a cantilever 23 titanium dioxide particles, gel swelling 23

476

Index

Vectors (polymeric nucleic acid delivery systems) 65–9 Vibrating actuators, see Thickness-mode actuation Vinylpyrrolidone 12 Viral gene therapy 65 Virtual reality tactile display (VRTD), Braille displays 440–1 Visual impairment, see Braille displays Void formation (crazing) 11 Volume changes, rate-limitation 8

Water loss under compression, gain under tension 9 self-diffusion coefficient 15 Word processing screen, typical 266 Young’s modulus 95, 213–14, 217 E-M polymers 326 radius of curvature 254