Surfaces and Interfaces for Biomaterials 1855739305, 9781855739307, 9781845690809

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Surfaces and interfaces for biomaterials

Related titles from Woodhead Publishing Limited: Medical textiles and biomaterials for healthcare (ISBN 1 85573 683 7) Medical textiles and biomaterials are a significant and increasingly important part of the technical textiles industry. They cover a huge range of applications from diapers and surgical gowns to substrates for electronic sensoring of vital life signs, external use as wound care, and internal use as implantables for biodegradable post-operative support systems, as well as the replacement of body parts through tissue engineering by supplying the structure for the growth of new cells. Even the humble plaster has the potential to deliver a powerful healthcare effect through its specific skincare characteristics and controlled delivery of medications. Medical textiles and biomaterials for healthcare will discuss recent developments in the main aspects of healthcare and medical textiles covering materials, manufacture, performance, applications, standards and user experience. It will serve as an essential resource for healthcare and medical textiles manufacturers, clinicians, researchers and consumers. Biomaterials, artificial organs and tissue engineering (ISBN 185573 737 X) Maintaining quality of life in an ageing population is one of the great challenges of the twenty-first century. This book and collection of illustrated CD lectures summarises how this challenge is being met by multi-disciplinary developments in specialty biomaterials, devices, artificial organs and in-vitro growth of human cells as tissue engineered constructs. Part A provides an introduction to living and man-made materials for the nonspecialist; Part B is an overview of clinical applications of various biomaterials and devices; Part C summarises the bioengineering principles, materials and designs used in artificial organs; Part D presents the concepts, cell techniques, scaffold materials and applications of tissue engineering; and Part E provides an overview of the complex socioeconomic factors involved in technology-based healthcare, including regulatory controls, technology transfer processes and ethical issues. Each chapter is supplemented with illustrated lectures and study questions in an easy to use CD to aid the reader in self-paced instruction. Hyaluronan (two-volume set: ISBN 1 85573 570 9) (Proceedings of an international meeting, September 2000, North East Wales Institute, UK.) Hyaluronan and its derivatives have developed very quickly in the last few years from a scientific novelty into an important new material for a diverse range of medical and biomaterial applications. The landmark conference on which this two-volume reference work is based focused on recent developments and applications in the use of hyaluronan in tissue repair and reconstruction, drug delivery systems, anti-cancer treatments and joint recovery and engineering. The entire range of hyaluronan progress is dealt with in depth by more than 135 individual papers presented in two volumes covering: analytical chemistry; chemical modification; physical characterisation; cell biology and medical applications. Details of these books and other Woodhead Publishing titles can be obtained by: · visiting our website at www.woodheadpublishing.com · contacting Customer Services (e-mail: [email protected]; fax: +44 (0) 1223 893694; tel.: +44 (0) 1223 891358 ext. 30; address: Woodhead Publishing Limited, Abington Hall, Abington, Cambridge CB1 6AH, England)

Surfaces and interfaces for biomaterials Edited by Pankaj Vadgama

Published by Woodhead Publishing Limited Abington Hall, Abington Cambridge CB1 6AH England www.woodheadpublishing.com Published in North America by CRC Press LLC 2000 Corporate Blvd, NW Boca Raton FL 33431 USA First published 2005, Woodhead Publishing Limited and CRC Press LLC ß 2005, Woodhead Publishing Limited The authors have asserted their moral rights. Every effort has been made to trace and acknowledge ownership of copyright. The publishers will be glad to hear from the copyright holders whom it has not been possible to contact concerning the following: Figs 1.11 and 1.13. This book contains information obtained from authentic and highly regarded sources. Reprinted material is quoted with permission, and sources are indicated. Reasonable efforts have been made to publish reliable data and information, but the authors and the publishers cannot assume responsibility for the validity of all materials. Neither the authors nor the publishers, nor anyone else associated with this publication, shall be liable for any loss, damage or liability directly or indirectly caused or alleged to be caused by this book. Neither this book nor any part may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, microfilming and recording, or by any information storage or retrieval system, without permission in writing from the publishers. The consent of Woodhead Publishing Limited and CRC Press LLC does not extend to copying for general distribution, for promotion, for creating new works, or for resale. Specific permission must be obtained in writing from Woodhead Publishing Limited or CRC Press LLC for such copying. Trademark notice: product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation, without intent to infringe. British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress Woodhead Publishing Limited ISBN-13: 978-1-85573-930-7 Woodhead Publishing Limited ISBN-10: 1-85573-930-5 CRC Press ISBN 0-8493-3446-6 CRC Press order number: WP3446 Project managed by Macfarlane Production Services, Markyate, Hertfordshire ([email protected]) Typeset by Godiva Publishing Services Ltd, Coventry, West Midlands Printed by TJ International Limited, Padstow, Cornwall, England

Contents

Contributor contact details

xv

Preface

xxi

Part I Forming methods 1

Fundamental properties of surfaces

3

P W E I G H T M A N and D S M A R T I N , The University of Liverpool, UK

1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9 1.10 1.11

2

Introduction Experimental considerations Surface characteristics Active sites and kinetics Controlling crystal growth: semiconductor technology Heterogeneous catalysis Real surfaces: theoretical advances Real surfaces: experimental approaches Insight into the biological activity of surfaces Conclusion References

Control of polymeric biomaterial surfaces

3 4 14 15 17 20 23 24 25 27 27

29

V H A S I R C I and N H A S I R C I , METU, Turkey 2.1 2.2 2.3 2.4 2.5 2.6

Introduction Preparation of polymers The solid state and structure Polymer-solvent interactions The polymeric surface and surface-bulk difference The general properties of a biomaterial surface

29 29 32 35 38 39

vi

Contents

2.7 2.8 2.9 2.10 2.11

Modification of polymer surfaces Surface analysis Surface properties and biomaterials applications Conclusion References

3

Organic thin film architectures: fabrication and properties

40 50 56 57 57

60

M C P E T T Y , University of Durham, UK 3.1 3.2 3.3 3.4 3.5 3.6 3.7

4

Introduction Established deposition methods Molecular architectures Molecular organization in thin films Future trends Further information References

Membranes and permeable films

60 61 65 71 78 79 79

83

N A H O E N I C H , University of Newcastle upon Tyne, UK and D M A L I K , Loughborough University, UK 4.1 4.2 4.3 4.4 4.5 4.6 4.7 4.8

5

Introduction 83 Materials and applications 83 Membrane characterisation 89 Blood material contact 92 Biological events at the membrane and thin film blood interface 93 Improvement of biocompatibility 99 Conclusion 99 References 100

Stable use of biosensors at the sample interface

103

J F G A R G I U L I , University of London, UK, A G I L L and G L I L L I E , University of Manchester, UK, M S C H O E N L E B E R , University of London, UK, J P E A R S O N , University of Manchester, UK, G K Y R I A K O U and P V A D G A M A , University of London, UK 5.1 5.2 5.3 5.4 5.5 5.6 5.7

Introduction Biosensor limitations Biocompatibility Materials interfacing strategy Membrane systems used in biosensors Microflows as surrogate, renewable barrier films Microfluidics and biosensors

103 104 105 116 119 138 140

Contents 5.8 5.9 5.10

6

Conclusion Acknowledgements References

Micro- and nanoscale surface patterning techniques for localising biomolecules and cells: the essence of nanobiotechnology

vii 146 147 147

150

Z A D E M O V I C and P K I N G S H O T T , The Danish Poymer Centre, Denmark

6.1 6.2 6.3 6.4 6.5 6.6 6.7

Introduction Lithographic patterning with photons, particles and scanning probes Soft lithographic techniques Colloidal-based fabrication techniques Template-imprinted nanostructured surfaces Conclusion References

150 152 158 168 169 169 171

Part II Measurement, monitoring and characterisation 7

Surface spectroscopies

183

M M E H L M A N N and G G A U G L I T Z , University of Tuebingen, Germany

7.1 7.2 7.3 7.4 7.5 7.6

8

Introduction Surfaces Optical detection methods Biomolecular interaction analysis Conclusion References

Surface microscopies

183 183 186 190 196 197

200

C Z I E G L E R , University of Kaiserslautern, and Institut fuÈr OberflaÈchen- und Schichtanalytik GmbH, Kaiserslautern, Germany

8.1 8.2 8.3 8.4 8.5 8.6 8.7

Introduction Electron microscopies Scanning probe microscopies Optical microscopies Future trends Further information References

200 208 211 218 219 220 220

viii

9

Contents

Nanoindentation

225

A B M A N N , The State University of New Jersey, USA 9.1 9.2 9.3 9.4 9.5 9.6 9.7 9.8 9.9

Introduction Instrumentation Data analysis Thin films and coatings Hard biological materials Soft biological materials Conclusion Further information References

10

Surface plasmon resonance

225 225 229 235 236 240 241 242 242

248

V H P EÂ R E Z - L U N A , Illinois Institute of Technology, USA 10.1 10.2 10.3 10.4 10.5 10.6 10.7

Introduction Surface plasmon resonance phenomenon Surface functionalization Applications Conclusion Acknowledgements References

11

Ellipsometry for optical surface study applications

248 249 257 261 264 265 265

271

Y M G E B R E M I C H A E L and K T V G R A T T A N , City University, London

11.1 11.2 11.3 11.4 11.5 11.6 11.7

12

Introduction History of ellipsometry and polarisation control Fibre based polarisation modulated ellipsometry A high birefringence fibre polarisation modulation ellipsometry Future trends Sources of further information References

Neutron reflection

271 278 282 285 292 292 294

299

J R L U , UMIST, UK 12.1 12.2 12.3 12.4 12.5 12.6

Introduction Neutron reflection and deuterium labelling Peptide interfacial assembly Lysozyme adsorption: the effect of surface chemistry Effect of size of globular proteins on their adsorption Conclusion

299 300 303 305 314 315

Contents 12.7 12.8

13

ix

Acknowledgements References

317 317

Microgravimetry

322

S V M I K H A L O V S K Y , University of Brighton, UK, V M G U N ' K O , Institute of Surface Chemistry, Ukraine, K D P A V E Y , University of Brighton, UK, P E T O M L I N S , National Physical Laboratory, UK and S L J A M E S , University of Brighton, UK

13.1 13.2 13.3 13.4 13.5 13.6 13.7 13.8 13.9 13.10 13.11 13.12

Introduction Quartz crystal microbalance technique Analytical applications of QCM Combination of QCM with other techniques Acoustic/piezoelectric sensors Future development of piezoelectric sensors Thermal gravimetry Non-QCM adsorption methods Dynamic contact angle measurements Conclusion Acknowledgements References

322 322 332 355 357 359 360 361 362 366 366 366

Part III Surface interaction and in-vitro studies 14

Interaction between biomaterials and cell tissues

389

Y I W A S A K I and N N A K A B A Y A S H I , Tokyo Medical and Dental University, Japan

14.1 14.2 14.3 14.4 14.5 14.6 14.7

15

Introduction Surface properties of biomedical materials Surface analyses of biomedical materials Design for non-biofouling surface How to connect tissues with biomaterials Conclusion References

389 389 393 399 405 410 411

Blood flow dynamics and surface interactions

414

W V A N O E V E R E N , University of Groningen, The Netherlands 15.1

Clinical application and problems of medical devices in contact with blood

414

x

Contents

15.2 15.3

Surface interactions of blood Role of blood cells during flow: rolling of cells, effect of concentration of erythrocytes, expression of adhesive cell receptors Biomaterial surface characteristics in relation to haemocompatibility and clinical applications Haemocompatibility of metals, ceramics and polymers Biological surface treatment to improve haemocompatibility ISO 10993 requirements for testing of medical devices: simulation of clinical application including flow, blood composition, anticoagulants Test models: static, low flow, arterial flow, pulsatile/laminar flow Conclusion References

15.4 15.5 15.6 15.7

15.8 15.9 15.10

16

Cell guidance through surface cues

418

422 423 425 427

431 432 434 435

447

A K V O G T - E I S E L E , Max-Planck Institute for Polymer Research, Germany, A O F F E N H AÈ U S S E R , Institute for Thin Films and Interfaces, Research Centre JuÈlich, Germany and W K N O L L , Max-Planck Institute for Polymer Research, Germany

16.1 16.2 16.3 16.4 16.5 16.6

17

Introduction Surface functionalization Patterning of chemical surface cues Synaptic connections in patterned neuronal networks: communication along predefined pathways Conclusion References

Controlled cell deposition techniques

447 451 454 458 461 462

465

C M A S O N , University College London, UK 17.1 17.2 17.3 17.4 17.5 17.6 17.7

Introduction In-vivo and in-vitro cell interactions Two-dimensional controlled cell deposition techniques Three-dimensional controlled cell deposition techniques Future trends Further information References

465 466 468 480 483 484 486

Contents

18

Biofouling in membrane separation systems

xi

493

Z C U I and Y W A N , University of Oxford, UK 18.1 18.2 18.3 18.4 18.5 18.6 18.7

Introduction Membrane separation ± concepts and applications Fouling mechanisms and factors affecting fouling Biofouling Fouling control Conclusion and future trends References

493 495 500 508 513 529 530

Part IV Surface interactions and in-vivo studies 19

Bioactive 3D scaffolds in regenerative medicine: the role of interface interactions

545

J R J O N E S and L L H E N C H , Imperial College London, UK 19.1 19.2 19.3 19.4 19.5 19.6 19.7 19.8 19.9 19.10

Introduction The need for biomedical materials and implants Surgical procedures for bone repair Surgical procedures in lung repair A new direction: regenerative medicine Bone regeneration Tissue engineering of the lung Conclusion Acknowledgements References

20

Intravascular drug delivery systems and devices: interactions at biointerface

545 545 547 550 551 552 561 567 568 568

573

K S R A O , Nebraska Medical Center, USA, A K P A N D A , National Institute of Immunology, India and

V L A B H A S E T W A R , Nebraska Medical Center, USA 20.1 20.2 20.3 20.4 20.5 20.6 20.7 20.8

Introduction Biomaterials and biointerface Intravascular drug delivery systems Nanoparticles as an intravascular delivery system Stents Vascular grafts and catheters Future trends References

573 573 575 575 580 581 581 582

xii

Contents

21

Surface degradation and microenvironmental outcomes

585

C C C H U , Cornell University, USA 21.1 21.2 21.3 21.4

585 587 589

21.5 21.6

Introduction Chemistry of synthetic biodegradable biomaterials In-vitro degradation of synthetic biodegradable biomaterials In-vivo biodegradation of synthetic biodegradable biomaterials and cell/biomaterial surface interaction Conclusion References

22

Microbial biofilms and clinical implants

619

601 614 614

M M I L L A R , Barts and the London School of Medicine and Dentistry 22.1 22.2 22.3 22.4 22.5 22.6 22.7 22.8 22.9 22.10 22.11

23

Introduction Epidemology and costs of infection associated with clinical implants Microbiology of clinical implant infections Molecular mechanisms underlying biofilm formation Determinants of biofilm antibiotic resistance Consequences of biofilm formation on clinical implants Clinical implant infection Prevention of biofilm formation on clinical implants Further research Information resources References

Extracellular matrix molecules in vascular tissue engineering

619 620 621 622 624 624 625 627 630 630 631

637

C M K I E L T Y , D V B A X , N H O D S O N and M J S H E R R A T T , Wellcome Trust Centre for Cell-Matrix Research, UK

23.1 23.2 23.3 23.4 23.5 23.6 23.7 23.8 23.9

Introduction Natural blood vessels Vascular tissue engineering Coating ECM molecules on surfaces ± a cautionary tale Biological seeding materials ECM-regulated delivery of therapeutic growth factors Future trends Acknowledgements References

637 638 641 646 651 657 658 658 659

Contents

24

Biomineralisation processes

xiii

666

F C M E L D R U M , University of Bristol, UK 24.1 24.2 24.3 24.4 24.5 24.6 24.7 24.8 24.9 24.10

25

Introduction `Biologically-induced' and `organic matrix-mediated' mineralisation Organic macromolecules Control over crystal structure Control over crystal orientation Control over morphology Control over mechanical properties Conclusion Further information References

666 667 667 671 674 676 686 687 688 688

On the topographical characterisation of biomaterial surfaces 693 P E T O M L I N S and R L E A C H , National Physical Laboratory, UK, P V A D G A M A , University of London, UK, S M I K H A L O V S K Y and S J A M E S , University of Brighton, UK

25.1 25.2 25.3 25.4 25.5 25.6 25.7 25.8 25.9 25.10 25.11 25.12 25.13 25.14

Introduction Biomaterials, surfaces and biocompatibility What is a surface? Surface measurement Filters Quantifying surface texture Two-dimensional profile data Three-dimensional data Techniques for surface texture measurement Traceability and calibration Conclusion Further reading Acknowledgements References

693 693 694 694 695 696 696 700 703 713 713 714 715 715

Part V Appendices 26

Surface modification of polymers to enhance biocompatibility

719

M T A V A K O L I , TWI Limited, UK 26.1 26.2

Introduction Polymers in medical applications

719 720

xiv

Contents

26.3 26.4 26.5 26.6 26.7

Biocompatibility Surface modification techniques Future trends Acknowledgements References

27

Issues concerning the use of assays of cell adhesion to biomaterials

722 723 740 741 741

745

S L J A M E S and S M I K H A L O V S K Y , University of Brighton, UK, P V A D G A M A , University of London, UK and P E T O M L I N S , National Physical Laboratory, UK 27.1 27.2 27.3 27.4 27.5 27.6 27.7

Introduction Measurement objectives Issues of interpretation of adhesion measurements Sources of variability in adhesion assays Methods of assaying cell adhesion in current use Conclusion References

28

Protein adsorption to surfaces and interfaces

745 746 750 753 757 760 761

763

B M I L T H O R P E , University of New South Wales, Australia 28.1 28.2 28.3 28.4 28.5 28.6 28.7 28.8

Introduction Classification of biomaterials surfaces and interfaces Non-specific adsorption to hard surfaces General rules of non-specific adsorption to flat surfaces Non-specific adsorption to `soft' surfaces Non-specific adsorption to penetrable surfaces and interfaces Future trends References

763 764 765 768 773 774 775 776

Index

782

Contributor contact details

Introduction Professor Pankaj Vadgama Director, IRC in Biomedical Materials Queen Mary University of London Mile End Road London E1 4NS UK E-mail: [email protected] Chapter 1 Professor Peter Weightman and Dr David S Martin Physics Department and Surface Science Research Centre The University of Liverpool Liverpool L69 3BX UK Tel: +44 (0)151 794 3871 E-mail: [email protected] E-mail: [email protected] Chapter 2 Professor V Hasirci Department of Biological Sciences Biotechnology Research Unit and Professor N Hasirci

Department of Chemistry METU Ankara 06531 Turkey E-mail: [email protected] E-mail: [email protected] Chapter 3 Professor M C Petty School of Engineering and Centre for Molecular and Nanoscale Electronics University of Durham South Road Durham DH1 3LE UK Tel: +44 (0)191 334 2419 E-mail: [email protected] Chapter 4 Dr Nicholas A Hoenich School of Clinical Medical Sciences Medical School University of Newcastle Framlington Place Newcastle upon Tyne NE2 4HH UK Tel: +44 (0)191 222 6998 E-mail: [email protected]

xvi

Contributor contact details

Dr Danish J Malik Department of Chemical Engineering Loughborough University Loughborough LE11 3TU UK Tel: +44 (0)1509 222507 E-mail: [email protected]

Chapter 5 Joseph Gargiuli IRC in Biomedical Materials Queen Mary, University of London Mile End Road London E1 4NS UK Tel: +44 (0)20 7882 5547/3254. E-mail: [email protected] Andrew Gill Inverness Medical Ltd Beechwood Business Park North Inverness IV2 3ED UK Tel: +44 (0)1463 721706. E-mail: [email protected] Monika Schoenleber IRC in Biomedical Materials Queen Mary, University of London Mile End Road London E1 4NS UK Tel: +44 (0)20 7882 3316 E-mail: [email protected] J Pearson Department of Epidemiology University of Manchester Manchester M13 9PL UK

G Kyriakou IRC in Biomedical Materials Queen Mary, University of London Mile End Road London E1 4NS UK Tel: +44 (0)20 7882 3316 E-mail: [email protected] Professor Pankaj Vadgama Director, IRC in Biomedical Materials Queen Mary, University of London Mile End Road London E1 4NS UK Tel: +44 (0)20 7882 3316 E-mail: [email protected] Chapter 6 Dr P Kingshott The Danish Polymer Centre Riso National Laboratory PO Box 49 Fredericksborgvej 399 DK-4000 Roskilde Denmark E-mail: [email protected] E-mail: [email protected] Chapter 7 Professor G Gauglitz and M Mehlmann Institut fuÈr Physikalische und Theoretische Chemie UniversitaÈt TuÈbingen Auf der Morgenstelle 9 72076 TuÈbingen Germany

Contributor contact details E-mail: [email protected] E-mail: [email protected]

Chapter 8 C Ziegler Department of Physics, University of Kaiserslautern and Institut fuÈr OberflaÈchen- und Schichtanalytik GmbH Erwin-SchroÈdinger-Straûe 56 67663 Kaiserslautern Germany Tel: +49 631 205 2855 E-mail: [email protected]

Chapter 9 Adrian Mann Department of Biomedical Engineering Rutgers The State University of New Jersey Piscataway NJ 08854 USA E-mail: [email protected]

Chapter 10 Professor Victor Perez Luna Room 144 Perlstein Hall Illinois Institute of Technology in Bioengineering, Chicago USA E-mail: [email protected]

xvii

Chapter 11 Dr Y M Gebremichael and Professor K T V Grattan School of Engineering and Mathematical Sciences City University Northampton Square London EC1V 0HB UK Tel: +44 (0)20 7040 3888 Fax: +44 (0)20 7040 8568 E-mail: [email protected] E-mail: [email protected]

Chapter 12 Professor Lu Biological Physics Group School of Physics and Astronomy The University of Manchester G13, Sackville Street Building Sackville Street Manchester M60 1QD UK Tel: +44 (0)161 2003926 E-mail: [email protected] Chapter 13 Professor S V Mikhalovsky, Dr K D Pavey and Dr S L James School of Pharmacy and Biomolecular Sciences, University of Brighton Cockcroft Building Lewes Road Brighton BN2 4GJ UK Tel: +44 (0)1273 642034 Fax: +44 (0)1273 642115 E-mail: [email protected]

xviii

Contributor contact details

Tel: +44 (0)1273 642042 Fax: +44 (0)1273 679333 E-mail: [email protected] Dr V M Gun'ko Institute of Surface Chemistry 17 General Naumov Street 03164 Kiev Ukraine Tel/Fax: +38 (044) 5123095 E-mail: [email protected] Dr P Tomlins Materials Centre, National Physical Laboratory, Queens Road, Teddington, Middlesex, TW11 0LW UK Tel: +44 (0)208 943 6778 Fax: +44 (0)208 943 6453 E-mail: [email protected] Chapter 14 Professor N Nakabayashi Institute of Biomaterials and Bioengineering Tokyo Medical and Dental University 2-3-10 Kanda-surugadai Chiyoda-ku Tokyo 101-0062 Japan Tel: +81-47-341-9734 E-mail:[email protected] Dr Yasuhiko Iwasaki Institute of Biomaterials and Bioengineering Tokyo Medical and Dental University

Kanda-surugadai Chiyoda-ku Tokyo 101-0026 Japan Tel: +81-3-5280-8026 E-mail: [email protected]

Chapter 15 Dr Willem van Oeveren, Dept of Biomedical Engineering, University Medical Center Groningen, Ant. Deusinglaan 1 PO Box 196, 9700AD Groningen The Netherlands Tel: +31 503633127 E-mail: [email protected]

Chapter 16 Professor W Knoll Max-Planck-Institut fuÈr Polymerforschung Postfach 3148 55021 Mainz Germany Tel: +49 06131 / 379 160 E-mail: [email protected] Dr Angela K Vogt-Eisele Cell Physiology ND/4 Ruhr University Bochum Universitystr. 150 44 780 Bochum Germany Tel: +49 (0)234 32 26756 Fax: +49 (0)234 32 14129 E-mail: [email protected]

Contributor contact details Professor Andreas OffenhaÈusser Institute for Thin Films and Interfaces (ISG-2) Forschungszentrum JuÈlich D-52425 JuÈlich Germany Tel: +49 (0)2461 612330 Fax: +49 (0)2461 618733 E-mail: [email protected] Chapter 17 Mr Chris Mason FRCS Department of Biochemical Engineering University College London Torrington Place London WC1E 7JE UK Tel: +44 (0)20 7679 0140 Fax: +44 (0)20 7209 0703 E-mail: [email protected] Chapter 18 Professor Zhanfeng Cui Department of Engineering Science Oxford University Parks Road Oxford OX1 3PJ UK Tel: +44 (0)1865 273118 E-mail: [email protected] Dr Yinhua Wan Department of Engineering Science Oxford University Parks Road Oxford OX1 3PJ UK Tel: +44 (0)1865 273059 E-mail: [email protected]

xix

Chapter 19 Professor Larry Hench and Dr Julian Jones Department of Materials Imperial College London South Kensington Campus London SW7 2AZ UK E-mail: [email protected] E-mail: [email protected] Chapter 20 Vinod Labhasetwar, Ph.D. College of Pharmacy Department of Pharmaceutical Sciences 986025 Nebraska Medical Center Omaha, NE 68198-6025 USA Tel: (402) 559-9021 E-mail: [email protected] Chapter 21 C Chu Department of Textiles and Apparel, and Biomedical Engineering Program Martha Van Rensselaer Hall Cornell University Ithaca, New York, 14853-4401 USA Tel: 607-255-1938 E-mail: [email protected] Chapter 22 Michael Millar Department of Microbiology Barts and the London School of Medicine and Dentistry Whitechapel

xx

Contributor contact details

London E1 2AD UK Tel: +44 (0)20 7377 7080 E-mail: [email protected]

University of London Mile End Road London E1 4NS UK E-mail: [email protected]

Chapter 23 Cay M Kielty, Daniel V Bax, Nigel Hodson and Michael J Sherratt School of Biological Sciences University of Manchester Michael Smith Building Oxford Road Manchester M13 9PT UK

Chapter 26 Professor Mehdi Tavakoli TWI Limited Granta Park Abington Cambridgeshire CB1 6AL UK E-mail: [email protected]

E-mail: [email protected] Chapter 24 Dr Fiona Meldrum School of Chemistry University of Bristol Cantock's Close Bristol BS8 1TS UK Tel: +44 (0)117 3317215 E-mail: [email protected] Chapter 25 Dr Paul Tomlins National Physical Laboratory Queens Road Teddington Middlesex TW11 0LW UK E-mail: [email protected] Professor Pankaj Vadgama Director, IRC in Biomedical Materials Queen Mary

Chapter 27 Dr S L James School of Pharmacy and Biomolecular Sciences University of Brighton Cockcroft Building Lewes Road Brighton BN2 4GJ UK Tel: +44 (0)1273 642042 Fax: +44 (0)1273 679333 E-mail: [email protected] Chapter 28 Professor B Milthorpe Head of School Graduate Scool of Biomedical Engineering University of New South Wales UNSW Sydney 2052 Australia Tel: +61-2-9385-3911 Fax: +61-2-9663-2108 E-mail: [email protected]

Preface

Biomaterials research has undergone a variety of evolutionary, one might say revolutionary, developments in recent years. The full entry of biomaterials as a credible modality in modern medicine was essentially triggered by the work of John Charnley on hip replacement. Then, as now, bulk materials properties and biomechanics took centre stage in view of the stringent mechanical and tribological demands of the implants. However, such issues cannot be the sole determinants of clinical outcome. Interest in bulk properties has inevitably shifted to the important consideration of the surface with the plethora of allied interfacial phenomena, conditioning clinical outcome. A consideration of the interface goes beyond the realms of an intellectual research exercise given that, even today, implants are better characterised by a bioincompatibility, rather than the hoped for `inertness' of biocompatibility. In mass terms, of course the surface is insignificant compared with the bulk. However, whilst the latter is a largely sequestered environment, amenable to at least some degree of engineering/physical science based prediction, the interface is the seat of some extraordinarily powerful, frontline biological responses. These variously conspire to attack the `non-self' implant firstly to try to degrade and eliminate it, and if that is not possible, to mask it off with an inert capsule; not a very promising outcome for precision engineering based therapeutics. The rules that govern the complex biological events are yet to be fully worked out, but it is a fair assumption that study of surface chemistry, organisation, topography, mechanics and physics, etc., will repay the effort. Unfortunately, the importance of surface processes in determining functional outcomes is not yet matched by the attention given in dedicated texts. The growing body of original literature also seems increasingly dissipated in analytical, chemical and biophysical journals. This book attempts to redress the information imbalance between surface and bulk descriptions or biomaterials. The series of chapters attempts to cover a spectrum from the fundamentals of surface structure and forming methods to biological and clinical outcomes. The link between experimental studies of surfaces in a controlled biomatrix and clinical results is a difficult one to make,

xxii

Preface

but at least the authoritative descriptions provided here will promote fresh thinking in this important area. Without measurement tools, understanding is restricted, so some key surface analytical techniques, both classical and developing (viz. probe microscopies, optics) are included. Ultimately, quantitative measurement is vital to understanding and so this book is in part an expression of the UK interest in metrology, represented through Department of Trade and Industry (DTI) programmes on materials metrology. I would like to express my sincere thanks to the many authors who were so patient with me and who contributed so effectively in making this book a viable exercise. My thanks also to Francis Dodds, Melanie Cotterell and their colleagues at Woodhead, whose continuous support and advice made the whole process both pleasurable and interesting and one that was so ably initiated by Gwen Jones. Finally my thanks to Catherine Jones of the IRC in Biomedical Materials for her unfailing assistance over the innumerable author interactions in modifying and adapting texts and for ensuring the project remained on track. Pankaj Vadgama IRC in Biomedical Materials University of London

Part I Forming methods

1 Fundamental properties of surfaces P W E I G H T M A N and D S M A R T I N , The University of Liverpool, UK

1.1

Introduction

Surface scientists occasionally observe that while God created solids, the Devil created surfaces. This phrase encapsulates the fact that while surfaces often dominate the behaviour of materials they are very difficult to study. Surfaces usually lack the high symmetry and purity of the interior of a solid and are often strongly influenced by adsorbed impurities from their environment. It is hard to overstate the significance of the Devil's work since surfaces are very important. Phenomena experienced in everyday life such as corrosion, adhesion, adsorption, friction and lubrication all occur at surfaces. More intimately, the crucial role played by surfaces in biocompatibility gives them an importance in the design of materials used in dentistry, contact lenses and medical implants such as hip joints and knee replacements. Industrial processes that occur at surfaces have a great impact on our lives and include crystal growth, semiconductor device manufacture and heterogeneous catalysis. Surface properties will also have a dominant influence on the emerging field of nanotechnology. The control of surface properties is thus essential to the function of a wide variety of materials. Clearly, then, a primary aim of surface scientists is to obtain a sufficient understanding of surfaces to make it possible to control surface properties. The field of surface science concerns fundamental, nanoscale investigations of surface phenomena that are both scientifically important and technologically relevant. The subject is relatively new since the experimental study of clean surfaces was delayed for many years by the crucial limitation that in a pressure in excess of 10ÿ6 mbar a clean surface adsorbs a monolayer of impurities within a few seconds. Thus experimental work on clean surfaces could not begin until the development in the 1960s of ultra-high vacuum (UHV) techniques, which reach pressures of 10ÿ10 mbar or less. This technological advance led to a rapid increase in surface studies and a proliferation of experimental techniques of which there is space here to give only the briefest of descriptions of some of the most important.

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The advances made in the understanding of surfaces in UHV conditions in the last forty years raise two questions. Firstly, to what extent are the surfaces that we now understand representative of the field in general? Secondly, does an understanding of the behaviour of surfaces in UHV provide a good guide to the behaviour of surfaces in the ambient conditions in which we find them in everyday life? The second of these questions is particularly pertinent to the role of surfaces in biomaterial applications and it will be addressed later in this chapter. The first question requires a rough estimate of how many surfaces are worthy of study since there are an infinite number of ways of terminating single crystals, to say nothing of the number of possible surfaces of amorphous materials. As a rough guide we note that Wyckoff's six volume classification1 lists ~7000 crystal structures. If we assume that each crystal structure has three important crystal faces and that it is appropriate to seek to understand the interactions that occur between each of these faces with the ten most important gases then we have a target for surface science of understanding ~200,000 surfaces. So far, the structures of ~1000 surfaces have been determined, about 1% of the target. We have a good understanding of the behaviour of the known surface structures in UHV and some understanding of the factors that are important in catalysis. In addition surface science has made major contributions to the considerable progress that has been made in the design and controlled growth of semiconductor systems. However, as will be made clear later, we are only just beginning to develop an understanding of surfaces in ambient conditions. The study of surfaces thus has a long way to go, particularly in addressing issues that are important in the real world.

1.2

Experimental considerations

We begin with a brief account of the importance of UHV and a description of single crystal surfaces. There are many experimental probes capable of detailed investigations of surfaces and interfaces, however, we have space to give only a brief overview of some of the most commonly used surface techniques. A more extensive introduction to these and many other techniques may be found in the books by Woodruff and Delchar,2 Zangwill,3 Prutton4 and Venables.5

1.2.1 Ultra-high vacuum As indicated in the introduction, the experimental investigation of the fundamental properties of surfaces have mostly taken place in UHV where the pressure is typically 10-10 mbar ± thirteen orders of magnitude lower than atmospheric pressure. A UHV environment is required to prepare a well-defined clean surface and maintain it for a sufficient time for experimental studies. In addition to surface preparation, a good vacuum is also a prerequisite for many of the experimental probes used to study surfaces since these probes are often

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5

based upon controlling the trajectories of electrons and ions. Vacuum technology has developed pumping systems capable of maintaining a UHV environment within stainless steel chambers for indefinite periods of time. Experiments are not exclusively performed on clean surfaces in UHV and the recent development of experimental probes that are capable of operating in a non-UHV environment is giving rise to an increasing trend of experimental studies of surfaces in ambient and liquid environments.

1.2.2 Crystal surfaces and surface preparation The majority of experimental surface studies have been performed on single crystals in order to simplify the atomic and electronic structure of the surface. Crystal surfaces can be prepared so as to consist of relatively large flat terraces made up of atoms of similar atomic coordination, with relatively few atoms associated with defect sites such as steps. The majority of single crystals grow in one of four basic structures: simple cubic (SC), face centred cubic (FCC), body centred cubic (BCC) or hexagonal close packed (HCP). When a single crystal is terminated by a surface, then, depending on the angle of the termination, different atomic arrangements are exposed. These different surfaces are described by the Miller indices and an introduction to this system of classification of crystal structures and surfaces can be found in ref. 6, which also explains the Wood notation that is used to describe the symmetry of surfaces. For FCC and BCC crystals, surface planes are defined by three integers. The three `low index' surfaces, (110), (111) and (100) created from the FCC and BCC structures are shown in Figs 1.1 and 1.2, respectively. The figures show that different crystal planes have different atomic densities and hence differences in free energy at the surface. The free energy of a surface is an important determinant of its behaviour as will be discussed later in considerations of crystal growth. Free energy consideration also influences the natural cleavage planes of crystalline materials and cleavage along a nonpreferred direction often results in a rough morphology composed of small areas of energetically preferred faces known as faceting. The cleavage of single crystals in UHV is one of the easiest ways of producing a clean surface. However, it can only be applied to the natural cleavage planes and these are not always the most important surfaces of a material. A variety of ways have been developed for producing clean surfaces on crystal faces that cannot be obtained by cleaving. These often involve bombarding the surface with argon ions in UHV to remove impurities followed by annealing to remove the structural damage. However, there is no cleaning procedure that works for all surfaces, and a considerable amount of effort is devoted to finding a recipe for the cleaning of a surface which is to be studied for the first time. Fundamental investigations of systems in ambient air are naturally limited to surfaces that can be prepared in air, and this draws attention to structural

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1.1 Surface structures created from cleavage of the FCC structure: (a) (110), (b) (111), and (c) (100). For the surface structures, the unit cells are shown: light spheres ˆ 1st layer atoms, dark spheres ˆ 2nd layer atoms.

stability in the presence of oxygen. Surfaces like graphite and mica that are stable against oxidation and which are easily cleaved in air to yield clean surfaces have naturally attracted attention. Cleaving graphite ± a layered material ± produces a well-defined surface consisting of large atomically flat terrace planes bounded by steps (Fig. 1.3). Such a relatively simple structure is ideal for investigating adsorption of organic molecules, and graphite has been used extensively for such studies.7,8 Gold surfaces are relatively inert to the effects of the atmosphere and this element provides some of the few single

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1.2 Surface structures created from cleavage of the BCC structure: (a) (110), (b) (111), and (c) (100). For the surface structures, the unit cells are shown: light spheres ˆ 1st layer atoms, dark spheres ˆ 2nd layer atoms, darkest spheres ˆ 3rd layer atoms.

crystal surfaces that can be prepared by the flame annealing technique for use as electrodes in electrochemical experiments at the solid/liquid interface. The surface structures shown in Figs 1.1 and 1.2 represent perfect terminations of the bulk crystal structure. The periodicity of the surface structures facilitates their representation by two-dimensional unit cells, which

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1.3 (a) Atomic structure of graphite showing layers (b) AFM image of the graphite surface showing atomically flat terraces and steps.

can be repeated to represent the extended surface structure. For a perfectly terminated crystal surface, the lengths of the surface unit cell are determined by the lengths of the unit vectors of the corresponding two-dimensional plane of the three-dimensional unit cell of the bulk crystal structure. This relationship is termed a (1  1) surface structure in the Wood terminology.6 Atoms at newly created surfaces sometimes adopt different positions from those expected from a perfect termination of the bulk crystal structure and such `reconstructed' surfaces may adopt a new symmetry with a periodicity that differs from that of the bulk crystal structure. This new periodicity can often be represented relative

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1.4 The FCC(110)-(1  2) reconstruction. Light spheres ˆ 1st layer atoms, dark spheres ˆ 2nd layer atoms, darkest spheres ˆ 3rd layer atoms.

to the symmetry of the unit cell of the bulk structure by a change in the ratio of the surface to bulk unit vectors. The surface periodicity is then described in terms of these ratios, n and m, as (n  m) in the Wood notation. Sometimes the representation of surface structures also requires a rotation of the axes of the surface unit cell relative to those of the bulk unit cell. Surface structures often involve a relaxation of the first few atomic planes in the direction normal to the surface. One example of a common reconstruction is the (1  2) `missing row' structure of the FCC(110) surface. In this reconstruction, every other ‰1 1 0Š row of atoms from the normal (1  1) surface is missing (Fig. 1.4) and this doubling of periodicity in one dimension is captured by the (1  2) Wood terminology. The clean (110) surfaces of Au, Pt and Ir at room temperature in UHV all adopt the (1  2) structure. There are many different surface reconstructions and surfaces often reconstruct as a result of the adsorption of foreign species. The surface reconstructions of the Si(100) and (111) surfaces are among the most important technologically and are shown in Figs 1.5 and 1.6. These reconstructions are driven by the reactivity of the unsatisfied covalent bonds of the Si surface atoms. The reconstruction of the Si(100) surface is easy to understand in that it arises from the pairing of adjacent Si surface atoms to form dimers. The dimers form rows and since the surface periodicity along the rows is the same as that of the bulk structure while the surface periodicity at right-angles to the dimer rows is twice that of the bulk structure this is termed a (1  2) reconstruction in the Wood notation. As Fig. 1.6 shows, the reconstruction that occurs on the Si(111) surface is very complex. It has a (7  7) periodicity relative to the bulk structure and its formation requires a considerable displacement and rearrangement of the atoms in the top three layers of the

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1.5 The Si(100) 1  2 reconstruction (courtesy Klaus Hermann, Fritz-HaberInstitut, Berlin).

Si(111) surface. The determination of the atomic positions in this structure took a considerable amount of time and required the application of all the major surface science experimental techniques. It is one of the many triumphs of the field. Once a well-defined single crystal surface is prepared, controlled experiments can then be performed such as investigating the first stages of oxidation or the interaction of the surface with molecular adsorbates. Structural defects such as adatoms, vacancies and steps, in addition to chemical impurities, can dramatically affect surface properties. It is often possible to control the introduction of these defects to surfaces and observe their impact on the system under investigation.

1.6 The Si(111) 7  7 reconstruction (a) STM image (Omicron Nanotechnology GmbH) and (b) surface structure model (reprinted from ref. 6 with permission of Oxford University Press).

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1.2.3 Determination of the elemental and chemical composition of surfaces: photoelectron and Auger electron spectroscopy Photoelectron spectroscopy involves illuminating a surface with photons of fixed energy and analysing the electrons subsequently emitted from the surface region. Atomic core hole states created by the X-ray emission of photoelectrons often decay by the Auger process which results in the emission of a second electron with an energy that is characteristic of the atomic element involved. These techniques can also be used to estimate the relative concentrations of the atomic species present in surfaces and the thickness of overlayers. Chemical bonds give rise to small changes in XPS and AES spectra that can be used to obtain insight into the chemical composition of surfaces. The electronic structure of a surface can be investigated using ultra-violet (UV) photoemission spectroscopy where the lower energy and momentum of UV photons relative to high energy X-rays allows the valence levels to be measured at high resolution. Unoccupied electronic bands can be examined using inverse photoemission spectroscopy (IPES).

1.2.4 Determination of surface structures: diffraction techniques The determination of surface structures on a macroscopic scale is obtained using diffraction techniques which, by their nature, emphasise periodicity. The more commonly used of these techniques are low-energy electron diffraction (LEED) and surface X-ray diffraction (SXRD). They give information on surface crystallography and can be used to determine the atomic structure of reconstructions and the overlayer structures of adsorbates. While LEED is restricted to a UHV environment, SXRD can be used to study surfaces at solid/ liquid interfaces.

1.2.5 Determination of the nature of molecular adsorption at surfaces: vibration spectroscopies and temperature programmed desorption Vibrational spectroscopies such as Raman spectroscopy and reflection absorption infra-red spectroscopy can provide information on molecular orientations at surfaces and on molecule-surface bonding. For example, whether a planar molecule adsorbed at a surface is flat-lying or upright can be distinguished. These techniques exploit the fact that chemical bonds absorb infra-red light at characteristic vibrational frequencies. An absorption spectrum reveals the chemical species present and shifts in frequencies from those expected from the `free' molecules in solution indicate interactions with the surface or other surface species.

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Temperature programmed desorption (TPD) is primarily used to determine the binding energies of molecules adsorbed at surfaces. In this technique a linearly increasing temperature gradient is applied to the sample and the molecular mass of desorbing species is recorded as a function of time. In this way, specific mass fragments can be identified and the strength of interactions and the activation energy for desorption can be determined since the least bound molecules will desorb first as the temperature is raised.

1.2.6 Techniques for the study of surfaces in ambient conditions: scanning probe and optical techniques Scanning probe techniques such as scanning tunnelling microscopy (STM) and atomic force microscopy (AFM) provide information on atomic and electronic structure on a local scale. They reveal the morphology of the surface and the degree and extent of surface roughness and they can achieve atomic resolution. The mechanism of image formation involves the surface of interest being brought into close contact with a sharp tip and exploiting the precise scanning motion that can be obtained in the plane of the surface by the use of piezoceramic crystals. The surface is then scanned underneath the tip whilst monitoring a certain property of interaction between the tip and surface. For STM, a tunnelling current, of the order of nA, between the tip and sample is monitored and maintained at a constant value by an electronic feedback loop. The measured current is extremely sensitive to tip-sample separation and increases as the separation decreases. Thus, to maintain a constant current, the sample ± due to its morphology ± is moved up or down perpendicular to the surface plane and this information is used to construct a three-dimensional image of the surface. In AFM, the tip is mounted at the free end of a flexible cantilever that is pivoted at the other end. A laser beam reflecting off the back of the cantilever monitors the deflection of the tip-cantilever arrangement by the morphology of the surface. A constant deflection is maintained by moving the surface up or down and an image sensitive to the height is recorded. The basic principles of the operation of STM and AFM are shown in Figs 1.7 and 1.8. Optical probes have the potential to study surfaces in ambient conditions but need to allow for the fact that only ~1% of the optical signal reflected from a solid arises from the surface. This dominance of the bulk contribution to the reflected signal is overcome in reflection anisotropy spectroscopy (RAS) and the related technique of reflection anisotropy microscopy (RAM) by measuring the difference in reflectance of normal incidence linearly polarised light between two orthogonal directions in the surface plane. When applied to a cubic crystal material, the contribution from the bulk cancels by symmetry and RAS becomes a surface-sensitive probe of optical anisotropy. RAS was developed to monitor semiconductor growth at high pressures9 and has developed into a probe of metal

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1.7 Basic schematic showing the principle of operation of an STM (courtesy Michael Schmid, Technische Universitaet, Wien).

1.8 Basic schematic showing the principle of operation of an AFM.

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surfaces10 with the potential to monitor molecular assembly at surfaces.11 Nonlinear optical techniques such as second harmonic generation (SHG) and sum frequency generation (SFG) are also surface sensitive and can be used to investigate chemical reactions at surfaces in a wide range of environments.12 After this brief introduction to the practical aspects of surface science we now review some of the most important conclusions reached from the study of surfaces.

1.3

Surface characteristics

Surfaces are usually expected to be flat, rigid and inert. However, this is rarely the case for real surfaces. At a microscopic level even the smoothest crystalline surface contains steps where atomic terraces of different height meet. Furthermore, these steps usually do not have straight edges but display kinks as they meander across a surface. The atoms at steps and kinks can have very different properties from atoms in the middle of terraces, and they often determine surface properties of technological importance such as the growth of crystals and the break up and assembly of molecules in catalytic reactions. Real surfaces are often rather flexible and are able to change their atomic structure in response to interactions with molecules.13,14 We have noted that the (110) surfaces of FCC metals often adopt a (1  2) missing row reconstruction (Fig. 1.4). This occurs because energetically the surface is finely balanced between the perfectly terminated and rather open (110) crystal plane and the (1  2) structure with its lower-energy close-packed (111) facets. Quite small changes to the surface electronic structure can induce changes in the structure of (110) surfaces of FCC crystals. In an electrochemical cell for example, the (1  1) to (1  2) phase transition can be induced on an Au(110) electrode by a change in charge density resulting from a change in the electrode potential and this change is reversible. Similar reconstructions can be induced by changing the potential on the other low index faces of Au in an electrochemical environment. In UHV the (1  2) reconstruction is the equilibrium structure of the Au(110) and Pt(110) faces though small amounts of adsorbed CO will change the Pt(110) surface back to a (1  1) structure. The opposite behaviour is found by the deposition of small amounts of alkali elements on the (110) faces of Ni, Cu, Pd and Ag which induce a (1  1) to (1  2) phase transition. The adsorption of larger molecules can induce more radical changes in surface structure. A dramatic example of this is the adsorption of hexa-tertbutyl-decacyclene (C60H66) on Cu(110).15 This molecule is approximately flat and it anchors to the Cu(110) surface through the formation of a characteristic trench in which 14 Cu atoms are dug out of the surface in a staggered arrangement involving two neighbouring close-packed rows, the free Cu atoms being added to kink sites at steps. When the molecules desorb, the surface is left with these characteristic trenches. It is as if the molecule leaves behind a footprint on the surface.

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Another remarkable example of adsorbate-induced restructuring is provided by the study by Bowker and co-workers14 on the adsorption of formic acid on a partially oxidised Cu(110) surface. The adsorption of oxygen creates a well known structure with (2  1) periodicity in which Cu atoms are incorporated into linear O-Cu-O-Cu rows orientated in [001] directions which are anchored at steps and which grow into islands of the (2  1) structure as the oxygen coverage increases. The formation of this structure is facilitated by the high mobility of Cu atoms at this surface. The adsorption of formic acid leads to a reaction with the (2  1) islands in which the Cu atoms at the end of the O-CuO-Cu units are released. These Cu atoms migrate and attach to more stable sites at the step edges between adjacent (2  1) islands. The subsequent desorption of formic acid leaves a surface in which the initial almost linear steps have been transformed into a sharp and jagged sawtooth structure.

1.4

Active sites and kinetics

The term `active site' refers to a specific atomic site on a surface where an interaction or reaction occurs. Active sites often involve atoms located at steps and kinks which have different coordination numbers from those of atoms bound in the terrace planes. The nature and influence of active sites can often be determined from TPD studies as illustrated by the following example of H2 adsorbed on Pt surfaces.13 The ideal Pt(111) surface is flat with hexagonal symmetry and one can identify three obvious adsorption sites as illustrated in Fig. 1.9; the top site where a molecule sits on top of an atom, the bridge site where a molecule is located between two atoms and the three-fold hollow site between three adjacent atoms. TPD studies of hydrogen desorption from the flat (111) surface show two overlapping peaks (Fig. 1.10) indicating that hydrogen occupies two sites on Pt(111) terraces with very similar binding energies. The Pt(557) surface consists of flat (111) terraces and steps aligned along a common direction giving a surface that is rather like a staircase (Fig. 1.9b). TPD results for this surface show peaks in the desorption of hydrogen that occur at two very different temperatures (Fig. 1.10). The peak at lower temperature corresponds to the peaks seen on the Pt(111) surface and is associated with adsorption on the (111) terraces and the peak at higher temperatures is associated with hydrogen that is more tightly bound to the higher coordination sites available at the steps. The TPD results for the desorption of hydrogen from the Pt(12,9,8) surface, which consists of (111) terraces and steps with the steps containing a large number of kink sites (Fig. 1.9c), is shown in Fig. 1.10. An additional peak is observed in the TPD results for this surface compared to the results for the Pt(557) surface (Fig. 1.10) and it is natural to associate this higher temperature peak with the binding of hydrogen to kink sites. Some insight into the nature of the adsorption of hydrogen on Pt surfaces can be obtained from TPD studies of surfaces which are dosed with a mixture of H2

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1.9 Surface structures of (a) Pt(111) showing the three possible adsorption sites, (b) Pt(557) and (c) Pt(12,9,8).

and D2. If the molecules dissociate on the surface it is possible that HD molecules will form prior to desorption. It is found that on well ordered Pt(111) surfaces the formation of HD is below the detection limits of the experiment indicating that the molecules do not dissociate on adsorption. However on stepped surfaces the production of HD indicates that all the molecules dissociate on adsorption. Clearly the symmetry and coordination of the adsorption site has an important influence on the details of the adsorption and desorption processes, a result that is borne out by a large number of studies on a wide variety of other systems.

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1.10 TPD results of H desorption from Pt(111), Pt(557) and Pt(12,9,8) (reprinted from ref. 13 with permission of John Wiley & Sons Inc.).

1.5

Controlling crystal growth: Semiconductor technology

One of the most impressive achievements of surface science has been in fostering the growth of semiconductor devices that in the last twenty years have revolutionised many aspects of technology particularly in the field of communication. This industry was not possible until the development of UHV technology and associated surface science techniques of surface preparation and characterisation. Indeed, all stages in the development of semiconductor devices have benefited from the exploitation of advances made in surface science. The key to this technology is establishing control over the growth of single crystal semiconductors and this is usually achieved by the molecular beam epitaxy (MBE) technique, which operates in a UHV environment, or by a variety of chemical vapour deposition (CVD) techniques which operate at higher pressures but in very controlled conditions. We can obtain some insight into the factors that are important in controlling crystal growth by considering the growth of crystals from solution in which the differences in the free energy of different crystal faces have a strong influence on the crystal shape. Provided there is sufficient kinetic freedom during the growth process then atoms being

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added to a crystal during growth will migrate to the face with the lowest free energy. This face will thus grow faster than other faces and, all other factors being equal, the equilibrium crystal shape will be the one in which the relative distances of each face from the nucleation centre is inversely proportional to the ratio of the free energies of the faces. Often, however, the `other factors' are not equal and crystal shape can be strongly influenced by factors such as temperature and the composition of the growth medium. This was discovered several hundred years ago when in 1783 Rome de l'Isle showed that octahedrons are formed instead of cubes if NaCl is grown in the presence of urine.16 This puzzling observation has recently been explained17 in terms of the role of the urea molecule in stabilising the (111) crystal face in solution during growth. A similar rebalancing of kinetic factors caused by subtle changes in the conditions of growth probably explains why natural diamonds tend to have octohedral shapes while those that are grown under high pressure and high-temperature conditions in the laboratory have a more cubic form though the morphology of the latter can be modified to a certain extent by varying the growth temperature. Diamond is also a good illustration of the fact that the structure of a solid is not always in equilibrium since the phase diagram of carbon (Fig. 1.11) shows that at normal pressures and temperatures the equilibrium structure of carbon is graphite, which consists of layers of strongly bonded carbon atoms (Fig. 1.3) with much weaker bonding between the layers. The cubic diamond form in which each carbon is at the centre of a tetrahedron of neighbouring carbon atoms (Fig. 1.12) is thus metastable at room temperature and pressure. The reason diamond does not readily transform into graphite is due to the energy barrier that

1.11 The phase diagram of carbon.

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1.12 The diamond crystal structure (reprinted from Introduction to Solid State Physics, 7th edition (1996) by C. Kittel with permission from John Wiley & Sons Inc.).

has to be overcome to re-orientate the inter-atomic bonds. The system is thus not able to reach equilibrium because kinetic factors prevent it from escaping from a local energy minimum. These kinetic factors can be exploited in the growth of diamond. The carbon phase diagram (Fig. 1.11) shows that it is reasonable to expect to be able to grow diamond at the high-pressure, high-temperature region on the border between the graphite and diamond phases. However, it is surprising that diamond can also be grown in the low-pressure and relatively low-temperature region by the chemical vapour deposition (CVD) process, which makes use of a microwave plasma excitation of a hydrocarbon vapour stream. The key to the CVD growth of diamond is to include a very low concentration of carbon in the feedstock gas: a typical growing medium consists of 99.5% H2 and 0.5% CH4. This ensures that the plasma is dominated by H ions and these ions terminate the surface carbon atoms of the diamond ensuring the maintenance of local tetrahedral and near tetrahedral bonding configurations. When an unsaturated surface carbon bond occurs it is unlikely that the adjacent carbon also has an unsaturated bond. Thus in these conditions the growth mode tends to promote the tetrahedral bonding of the diamond structure rather than the planar geometry of the graphite structure. Clearly, in the manufacture of semiconductor devices, it is important to monitor and control surface processes and in particular to achieve an understanding of the kinetic factors that can be exploited to ensure that a complex multilayer device structure maintains its coherence. This control is achieved by the MBE and CVD techniques. We comment here on one small aspect of this field that illustrates the importance of an understanding of surface properties in the design and production of semiconductor devices. A key material of course is Si and at certain stages in device processing it is necessary to `passivate' Si surfaces so that they do not form chemical bonds with molecules in the growth environment. However Si surfaces are very reactive

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because Si adopts the diamond structure in which each atom bonds with four neighbouring atoms (Fig. 1.12) and when a surface is created this leaves the surface atoms with unsatisfied bonds. It is often necessary to passivate such surfaces at various stages in device manufacture. One approach is to deposit an ordered layer of As atoms taking advantage of the natural tendency of As to bond to three neighbours. The difference in the geometry of the Si(100) and (111) surfaces requires different passivating structures. The Si(100) surface is passivated by the formation of As dimers each of which bonds to two Si atoms in the layer below and its neighbouring As in the dimer. On the (111) surface As atoms are coordinated to three Si atoms in the layer below. In each of these structures the As atoms have no unsaturated bonds and the surface is inert ± as illustrated by the fact that the sticking coefficient for the reaction with oxygen is reduced by ~15 orders of magnitude with respect to the corresponding Si surface. This control of surface reactivity is just one of the many ways in which semiconductor device technology is dependent on advances made in surface science.

1.6

Heterogeneous catalysis

It is estimated that industrial processes that employ catalysts have a turnover of the order of £2,000 billion per year in the UK and quite small improvements in the efficiency of catalytic processes can have a major impact on industrial costs. When this is set against a background in which little is known about the actual mechanisms of catalysis one can understand why this field is one of the strongest drivers for research in surface science. The industrial exploitation of catalysis has already had a major impact on the world and one process in particular, the Haber-Bosch method of ammonia synthesis from a mixture of hydrogen and nitrogen gas, has been called the most important invention of the twentieth century because of its role in sustaining the increase in world population through the production of nitrogen based fertiliser (Fig. 1.13).18 This process was discovered empirically and was first used commercially in 1913. The catalytic mechanism is very complex and in spite of a century of research it is still not fully understood. The Haber-Bosch process involves passing N2 and H2 over Fe surfaces. The catalyst is composed of small porous Fe particles of surface area approximately 15 m2/g. A few percent of aluminium oxide (Al2O3), potassium oxide (K2O) and other compounds are added as `promoters' which increase the activity of the iron though their precise role is unclear. The reaction takes place at about 400 ëC in a total pressure of 150 to 300 atmospheres. Ammonia production shows a complex dependence on the partial pressures of N2, H2 and NH3. The reaction can be summarised as follows: N2 and H2 molecules interact with the surface, dissociate into atoms and chemisorb on the surface. The N atoms then react with H atoms to form NH3, which then desorbs from the surface. The main factors

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1.13 Growth in global consumption of N fertiliser and increase in world population (figure by Bryan Christie, reprinted from ref. 18).

that determine the rate of ammonia synthesis are the N2 dissociative sticking probability and the N atom chemisorption energy. As illustrated by the Haber-Bosch process a catalyst must be able to bind molecules, often dissociating them in the process, allow the bound species to react with other molecules and then allow desorption of the desired reaction product. Heterogeneous catalyst systems usually involve small particles that are dispersed over a support material ± often a metal oxide. The small particles have size of the order of nanometres leading to the term `nanoparticles'. Surface effects dominate the structural features of nanoparticles, which have high free energies and are thus very reactive. The fact that catalytic processes usually take place at high pressures means that there is a `pressure gap' in trying to obtain insight into catalytic processes from experiments in UHV conditions. Another limitation on the applicability of surface science techniques to the study of catalysis is that industrial catalysts usually employ highly reactive and poorly characterised nanoparticles whereas surface science seeks to simplify the complexities of chemical reactions by

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experimental work on well characterised single crystal surfaces. These limitations aside, surface science has provided some insight into catalytic mechanisms.19 As an illustration of the strengths and limitations of the current level of understanding achieved in surface science we consider two important examples: the Haber-Bosch process of ammonia synthesis and the oxidation of CO in catalysts for automobile exhausts.

1.6.1 The Haber-Bosch process Surface science studies of single crystal surfaces in UHV have yielded some insight into the mechanisms of the Haber-Bosch process. The results are summarised in ref. 13 and experiments on different surface planes of Fe single crystals show that the efficiency of the process reduces in the sequence: (111), (211), (100), (210), (110). These surfaces have different morphologies, and close examination shows that the (111) surface of BCC Fe is relatively rough exposing the first, second and third layer atoms while the (110) surface is relatively smooth exposing the first layer of atoms and a fraction of the second layer (Fig. 1.2). However surface roughness alone cannot explain the observed catalytic activities which are actually determined by the nature of the active sites. High coordination sites, as found on (111) and (211) surfaces, promote reactions and a good correlation has been found between coordination and efficiency. This suggests why small particles are more efficient than larger ones since small particles have larger numbers of high coordination sites.

1.6.2 The oxidation of CO in automobile exhausts In recent years catalysis has found increasing application in mitigating the environmental impact of automobile exhausts. A typical catalytic converter in a car exhaust is constructed from a high surface area honeycomb structure which is coated with a thin mixture of porous aluminium oxide, ceria and zirconia that is impregnated with nanoparticles of reactive metals such as platinum, ruthenium, palladium and rhodium. Platinum is used to oxidise hydrocarbons and carbon monoxide gas while rhodium serves to reduce nitrogen-oxide species. Early work seeking to understand the detail of these reactions adopted the approach of characterising the surface in UHV, conducting a reaction at high pressure and then examining the surface immediately afterwards. Important information has been gained by this approach, revealing changes in structure and composition of the surface following the catalytic reaction. However, the need to monitor and understand such processes at high pressure is illustrated dramatically by the observation that whereas Ru is the least efficient metal for promoting the oxidation of CO in UHV, it is the most efficient for promoting this reaction at high pressures. Fortunately, significant progress is now being made both theoretically and experimentally in this field and the resolution of the

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apparent paradox of the pressure dependence of the oxidation of CO by Ru, which is described below, is an excellent illustration of the importance of bridging the pressure gap in the study of surfaces.

1.7

Real surfaces: theoretical advances

In spite of the progress made in surface science in the last forty years, our understanding of the behaviour of surfaces in the real world at normal temperatures and pressures is currently quite poor. However, advances are now being made in the study of real surfaces and much of the impetus for this progress has come from theoretical developments. It is now understood why the capacity of Ru to promote the oxidation of CO is so dependent on pressure. The insights into the resolution of this problem are likely to be generally relevant and are worthy of a brief summary. The key issue is to note that the length and timescales at which processes happen and can be studied experimentally in surface science are very different from those that determine the behaviour of surfaces in the real world. This issue has been reviewed recently by Stampf and co-workers20 who draw attention to three important regiems of length and timescale as shown in Fig. 1.14. The bottom left, microscopic, region of Fig. 1.14 is dominated by the motion of electrons and atoms which occur on timescales of 10ÿ15 seconds and 10ÿ12 seconds respectively. An appropriate length scale is 0.1 nm ± the length over which an atom moves between adjacent atomic sites. This regiem is well described by density functional theory which describes the behaviour of

1.14 The spatial and temporal domains relevant to different aspects of the behaviour of surfaces (reprinted from ref. 20 with permission from Elsevier).

24

Surfaces and interfaces for biomaterials

electrons in the potential environment created by atoms and molecules. Many surface science techniques yield useful information on processes that occur on the length and timescales that characterise this regiem. However, while the behaviour of systems in the microscopic regiem is a determinant of what can happen in the mesoscopic regiem, it is not possible to use results obtained in this regiem to infer what will happen at more macroscopic length and timescales. This is because an understanding of the more macroscopic regiems requires the introduction of approaches from statistical mechanics and thermodynamics. Fortunately, theoretical work in association with relevant experimental results is beginning to make useful links between regiems shown in Fig. 1.14. This work resolves the paradox concerning the behaviour of the Ru surface at high pressure since it makes clear that at high pressure the Ru surface converts to the oxide RuO2 and it is the oxide surface that promotes the oxidation of CO to CO2. The key observation is that at higher temperatures and pressures changes in the relative free energies of the metal and its oxide give rise to an intimate interplay between the progress of the chemical reaction and the morphology and chemical composition of the surface and the surrounding gases. We might expect similar considerations to be important in many other areas of surface science where work in UHV does not lead to a simple understanding of the behaviour of surfaces in the real world.

1.8

Real surfaces: experimental approaches

It is clear from the previous section that an understanding of the behaviour of real surfaces under ambient conditions requires information that can only be obtained using experimental probes that can operate outside of UHV. We saw earlier that scanning probe microscopy and optical probes both have this capability. Since, in general, scanning probes reveal information on the microand nanoscale and are rather slow and optical probes yield an average signal over a macroscopic area and can be rather fast these generic techniques complement each other in both the length and timescale over which they can provide information on surface processes. A good example is the study of Hendriksen and Frenken21 who used STM to observe in situ the oxidation of CO on the Pt(110) surface at semi-realistic conditions of sample temperature 425 K and gas flow of pressure 500 mbar. During the reaction the surface was found to change from a metallic CO-covered surface to an oxide surface. In a similar finding to the Ru case, the Pt-oxide surface was found to exhibit the higher catalytic activity. The potential of optical probes to reveal the dynamics of surface processes at higher pressures is shown by the studies of Rotermund and co-workers22 who used RAM to monitor the oxidation of CO at the Pt(110) surface in both low, ~10ÿ4 mbar, and high, approaching atmospheric, partial pressures of CO and O2. This surface has two states, an inert passive state in which the surface is

Fundamental properties of surfaces

25

`poisoned' by the presence of adsorbed CO and an active oxidising surface covered by the dissociative adsorption of O2. Since the surface is not perfect there will always be a particularly active site which at low pressure will promote local oxidation of CO to CO2. This reaction proceeds exothermically and the energy released locally initiates a reaction-front which, in the low partial pressure regimen, spreads across the surface leaving in its wake the poisoned surface created by the exhaustion of locally adsorbed oxygen. The progress of the reaction-front can be observed in real time by RAM and this technique shows that at relatively low pressures, the surface undergoes an oscillating reaction mediated by variations in the partial pressures of O2, CO and CO2. The oscillations are sufficiently robust that two crystal surfaces exposed to the same gaseous environment will oscillate in phase. It is important to note that in the low pressure regimen the thermal characteristics of these reactions are dominated by the large thermal capacity of the experimental mount that is supporting the Pt crystal. However, as the pressure of the reacting gases is increased, a point is reached where energy is released at the reaction site faster than it can be conducted away. The resulting local increase in temperature accelerates the reaction resulting in a thermal runaway that is terminated only by the local exhaustion of the reacting species. In these conditions the smoothly varying oscillations are replaced by sudden local bursts of activity that have no spatial or temporal coordination. This later mode of activity is likely to be more characteristic of the actual behaviour occurring during catalytic reactions at high pressure and demonstrates the importance of experimental work that can bridge the pressure gap.

1.9

Insight into the biological activity of surfaces

This brief survey of progress in surface science suggests a number of ways in which the characteristics of surfaces are likely to have biological significance. We highlight three areas.

1.9.1 The biological activity of nanoparticles As indicated, nanoparticulate matter is characterised by high free energies and a high concentration of active sites. For these reasons they find application in industrial catalytic processes. However, we should expect that these characteristics also endow nanoparticles with significant biological activity and it is well known that ultra-fine air-borne particulates can have harmful effects, as recently demonstrated by the studies of the damage they cause to the mitochondria in cells.23 As the emerging field of nanotechnology develops it would be sensible to be aware of the implications for the health of the population by the release of novel and biologically active particulates into the environment.24

26

Surfaces and interfaces for biomaterials

1.9.2 Insights into naturally occurring biological processes We have drawn attention to the enormous impact on the world of the application of the Haber-Bosch process for fixing nitrogen and the insight that surface science is providing into the detailed mechanisms of this reaction. However, in comparison with natural mechanisms of fixing nitrogen, the Haber-Bosch process is remarkably inefficient requiring, as it does, high pressures and high temperatures. Natural processes clearly extract nitrogen from the atmosphere and bind it in chemical complexes at normal temperature and pressure. Can surface science provide insight into the mechanisms of natural processes in fixing nitrogen? It is interesting to view the recent advance in the understanding of how Mo complexes fix nitrogen25 in a surface science context. To begin with, the mechanism requires a large molecular complex of nanoparticle dimensions. It also depends on an active catalytic site that shows considerable flexibility during the various reactions that constitute the mechanism of nitrogen fixation. This is consistent with what we have learnt from surface science concerning the importance of free energy, active sites and surface flexibility in promoting chemical reactions. One may speculate that the ultimate success of surface science will be to provide us with a sufficient understanding of catalytic processes that we can escape the restrictions imposed by surfaces and develop ways of constructing tailored active and flexible sites in molecules similar to the complexes that operate in natural processes.

1.9.3 Prospects for the study of biomedical interfaces Surface science has addressed metal, oxide and semiconductor surfaces with an approach involving simple systems studied at the fundamental level with powerful probes. This approach will be beneficial for the investigation of complex biosurfaces. One would anticipate that the important determinants of surface processes such as free energies, kinetic factors and the nature of active sites will have an important influence on the behaviour of biomedical interfaces. Clearly the study of such interfaces would benefit considerably from the development of experimental techniques for studying surfaces in ambient conditions. Some indications of the insight into the behaviour of biomedical interfaces can be obtained from studies of the temporal dependence of proteins deposited on human colostrum immobilised on methylated silicon surfaces from human blood serum.26 This study employed a combination of ellipsometry and antibody techniques to provide a convenient and rapid way to indicate the activation of complement sequences on solid surfaces and facilitated a time-resolved determination of replacement sequences and activation pathways.

Fundamental properties of surfaces

27

1.10 Conclusion In this chapter we have considered some of the fundamental properties of surfaces. We have seen that surface science has had a huge impact on the growth of semiconductor devices and has had some success in the understanding of catalytic processes. We have seen that progress is now being made, both theoretically and experimentally, in the study of real surfaces in ambient conditions. We believe that the insights obtained from the study of surfaces have the potential to make a major contribution to the understanding of biological systems.

1.11 References 1. Wyckoff R W G, Crystal Structures, 2nd edn, Volumes 1±6, Interscience, New York, 1963±1971. 2. Woodruff D P and Delchar T A, in Modern Techniques of Surface Science, 2nd edn, Cambridge University Press, 1994. 3. Zangwill A, Physics at Surfaces, Cambridge University Press, 1996. 4. Prutton M, in Introduction to Surface Physics, Clarendon Press, 1994. 5. Venables J A, in Introduction to Surface and Thin Film Processes, Cambridge University Press, 2000. 6. Attard G and Barnes C, in Surfaces, Oxford University Press 1998. 7. Magonov S N and Whangbo M-H, in Surface Analysis with STM and AFM, VCH, Weinheim, 1996. 8. Martin D S and Weightman P, Surf. Sci. 464 23 (2000). 9. Aspnes D E, Harbison J P, Studna A A and Florez L T, J. Vac. Sci. Technol. A6 1327 (1988). 10. Martin D S and Weightman P, Surface Review and Letters, 7 4 389 (2000). 11. Martin D S and Weightman P, Thin Solid Films, 455±456, 752 (2004). 12. McGilp J F, Surf. Rev. Lett., 6 (3±4), 529 (1999). 13. Somorjai G A, in Introduction to Surface Chemistry and Catalysis, Wiley, New York, 1994. 14. Bowker M and Bennet R A, Topics in Catalysis, 14 1 (2001). 15. Schunack M, Petersen L, Kuehnle A, Laegsgaard E, Stensgaard I, Johannsen I and Besenbacher F, Phys. Rev. Lett., 86 456 (2001). 16. de Rome de l'Isle J B L, Crystallographie (Imprimerie de Monsieur, Paris) 1 379 (1783). 17. Radenovic N, von Enckevoly W, Verwer P and Vlieg E, Surf. Sci., 523 307 (2003). 18. Smil V, `Global population and the nitrogen cycle', Scientific American, 227 76, July (1997). 19. Sinfelt J H, Surf. Sci., 500 923 (2002). 20. Stampfl C, Ganduglia-Pirovano M V, Reuter K and Scheffler M, Surf. Sci., 500 368 (2002). 21. Hendriksen B L M and Frenken J M W, Phys. Rev. Lett., 89, 46101 (2002). 22. Dicke J, Erichesen P, Wolff J and Rotermund H H, Surf. Sci., 462 90 (2000). 23. Li N, Sioutas C, Cho A, Schmitz D, Misra C, Sempf J, Wang M, Oberley T, Froines

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Surfaces and interfaces for biomaterials

J and Nel A, Environmental Health Perspectives, 111 455 (2003). 24. Report of Royal Society and Royal Academy of Engineering on Nanotechnology, July 2004. 25. Yandulov D V and Schrock R R, Science, 301 76 (2003). 26. Tengvall P, Askendal A and Lundstrom I, J. Biomedical Materials Research, 35 81 (1997).

2 Control of polymeric biomaterial surfaces V H A S I R C I and N H A S I R C I , METU, Turkey

2.1

Introduction

Polymers are long, linear or branched chains of molecular weight 10,000 daltons or higher consisting of one or more basic structural or monomeric units. The flexibility of the chain depends on the stiffness of the individual units making up the chain, to the bonds that link them together, the length of the chain and the presence and nature of the branches. These chains are found naturally both in the environment and in the human body. They can also be synthesized from the monomers in the laboratory to yield macromolecules of the desired properties.

2.2

Preparation of polymers

Polymers are prepared by the reaction of monomers. The method of polymerization could be classified as `condensation polymerization' and `addition polymerization'. Condensation polymers are formed by molecules with functional groups that could condense to form the chain mostly by the release of a small molecule such as water, ammonia, etc. Addition polymerization require unsaturated molecules to become saturated while adding to the chain after activation by ionic or free radical initiators, or high energy radiation. The synthetic process is started by the use of an initiator and is helped by exposure to highly energetic radiation (i.e. UV, gamma), elevated temperatures and accelerators. Biological polymers (such as proteins and polysaccharides) are formed exclusively through polycondensation but require enzymes for the initiation and progression of the chain formation reaction.

2.2.1 Homopolymerization Polymers could form by the chain reaction of a single monomer through a condensation or addition reaction. The resultant compound is a very large

30

Surfaces and interfaces for biomaterials

molecule consisting of a single molecule (could be represented by A below) repeated hundreds or thousands of times. -A-A-A-A-A-A-A-A-A-A-A-A-A-A-A-

or

-[A]n-

This kind of a molecule is generally linear unless higher functionality is available on the monomer. Some biological macromolecules such as polysaccharides could yield branched structures only because of the high number of -OH groups. Actually there are no natural homopolymers consisting of amino acids or nucleic acids to yield proteins (i.e. polyalanine) or DNA or RNA (i.e. polyU) from a single molecule. A typical example of this kind of polymer is polyethylene which consists of many -CH2-CH2- groups added on the end. This group is called the repeating unit, since it is the basic element of the macromolecule that repeats itself to represent the final product. The number of these repeating units is known as the degree of polymerization, P, which is an average value due to the random nature of the progression of the reaction. When the carbon atoms in the chain are asymmetric (like -CHR- and unlike -CH2-) the steric position of the monomer side group (R) gains importance. An irregular organization of the R groups lead to an atactic polymer while an alternating position leads to a syndiotactic polymer and a same side orientation is called isotactic. Crystallinity of polymers, a very important property of the solid state, is a direct result of this preferential orientation of the side groups.

2.2.2 Copolymerization and terpolymerization (alternating, block, random) Polymers can also form when two or more monomers are involved in the preparation of the macromolecule. The order in which these monomers come together is very influential on the properties of the resultant polymer. If there are two different monomers such as A and B they could take places in the chain in a random manner forming a random copolymer of A and B: -A-A-B-A-B-B-A-B-B-B-A-A-B-A-A-AThis is the case with most industrial copolymers where the composition of the monomer mixture is not reflected in the resultant polymer. The structural units of the polymer could be ordered in an alternating manner to yield an alternating polymer: -A-B-A-B-A-B-A-B-A-B-A-B-A-B-A-BAnother way the monomer could be arranged on the chain is by way of combination of segments or blocks of single monomer chains: A-A-A-A-A-B-B-B-B-B-B-A-A-A-A-A-A-

Control of polymeric biomaterial surfaces

31

2.2.3 Branching and crosslinking In principle, monomers are either bifunctional or they have unsaturations which enable them to interact with two molecules. In either case a linear molecule is formed. In the presence of compounds with more than two functional groups (for condensation) or more than one unsaturation (for addition) a branched molecule is formed. The branches might form simultaneously with the formation of the chains but also could be induced after the polymer formation is completed (Fig. 2.1). Crosslinking is similar to branching but the general understanding is that a number of chains are bonded together with groups which might (or might not) be of the same chemical origin (Fig. 2.2). As the number of crosslinks increase the solubility of the polymeric structure decreases. At a critical number all the chains of the polymer are bonded covalently together leading to an insoluble polymer, the fully crosslinked structure.

2.1 Branched polymers are linear polymers with short chains of the same composition attached.

2.2 Crosslinking involves covalent and transient linking of polymer chains and significantly modifies the mechanical and solution properties.

2.2.4 Grafting and blending In addition to the above there are graft polymers where one chain serves as a backbone and relatively shorter chains of different chemistry are attached to this backbone as branches (Fig. 2.3). Blending, on the other hand, is a physical mixture where there is no covalent bond between chains of two or more different types brought together either by melting or dissolution followed by solvent removal leading to macromolecules of various types of chains.

32

Surfaces and interfaces for biomaterials

2.3 Grafting of B oligomers on the poly A backbone.

2.3

The solid state and structure

Synthetic polymers are formed in the reaction media and each chain is initiated and terminated at a different time and as a result each chain has a different length leading to a distribution of chain lengths. This is unlike the biological system where the genetic material and proteins are of known fixed molecular weights because their production is highly controlled. In the chemical system there is no such control. As a result, rather than a single molecular weight, one determines an average molecular weight, the value of which is dependent on the method employed. The most widely used average molecular weights are weight average molecular weight (Mw ) and number average molecular weight (Mn ). Their ratio is called the Heterogeneity Index (HI) or polydispersity: w M HI ˆ  Mn and as the HI gets closer to 1 the molecular weight distribution gets narrower. This ratio is especially important in terms of the mechanical properties because as the fraction of very low molecular weight chains increase they act more as plasticizers that reduce the mechanical strength of the material. It is also an important indicator for the mobility of low molecular weight materials to the surface when the polymeric product is introduced into solvents.

2.3.1 Amorphous and crystalline polymers A glassy polymer structure is visualized as densely packed, entangled, random Gaussian coils (Fig. 2.4). The random coil state remains favourable in the glassy state (Stachurski, 2003). Above the glass transition temperature, Tg, where chains gain mobility, the chains are confined to molecular tubes, within which de Gennes reptation takes place. From NMR studies it is confirmed that the chain is not a simple (smooth), constant curvature random coil, but it is further segmented by folds (on itself). The number of folds per chain is a property of the chain (stiffness and chemistry) and of the temperature. The size of the fold is, to a good

Control of polymeric biomaterial surfaces

33

2.4 A glassy polymer structure.

approximation, independent of temperature and chain length. The segments between the folds are between 10 and 20 covalent bonds long. Within the macroscopic volume of the polymer, so-called `free volume' is present as an equilibrium property of the system at Tg. At sufficiently fast cooling rate from this state to a temperature below Tg, the polymer retains a certain amount of the free volume. The actual amount depends on the temperature from which the quench occurs, the cooling rate, and the type of the polymer and its volume. A glassy polymer is formed because the irregular chain architecture prevents crystallization. On cooling to below Tg, a portion of the unoccupied free volume spontaneously diffuses out, allowing the bulk volume to reduce. In the process, the equilibrium end-to-end distance of chains is compressed. Its relaxation time rapidly increases. The prerequisites for the ability to crystallize is regularity, both in terms of composition and stereochemistry of the polymer. Even then it is not enough for the observation of crystallinity. Long periods of cooling might be necessary for allowing the chain substituents to take the required conformation before they can crystallize. They can not, however, completely crystallize; there are always regions (chain ends or side groups) found not to be crystallized when examined by electron microscopy. The resultant structure is what is called the fringed micelle (Fig. 2.5). The rate of cooling is also important because if the cooling rate is high the polymer might not find time to orientate and crystallize and instead solidify in a glassy form. Polymer molecules in the melt are in the form of coils rather than stretched out chains. During the formation of the crystals the coil is still the Ê and the central form. A typical single crystal lamella thickness is around 100 A polymers arrange perpendicular to the plane of the crystal. This has been shown to be true with polyethylene, polypropylene, and low molecular weight paraffins. Since, however, the chains in the lamellae are not equal in length the surface of the single crystal must appear rough.

34

Surfaces and interfaces for biomaterials

2.5 Fringed micelle structure.

2.3.2 Packing, packing density and free volume A glassy polymer is formed because its irregular chain architecture prevents crystallization and it solidifies with a significant volume unoccupied by atoms, which is called the free volume. If this cooling rate is decreased then better packing with a lower free volume is observed. On the other hand when a crystalline polymer is solidified, the organization of the chains and the atoms are such that the unoccupied volume within the solid is minimal. Thus, the free volume in the amorphous polymer is higher than that in a crystalline polymer. The lower free volume indicates a higher amount of interaction between neighboring chain segments and thus to a mechanically stronger structure. In a fringed micelle structure where there is partial crystallinity, certain regions of the polymer are densely packed with low free volume, while the rest is amorphous with high free volume. In such a case the bulk of the polymer as well as its surface has regions of high and low packing density and is expected to have different responses in a biological medium.

2.3.3 Phase states and phase transition Polymers in a solid state could show different properties depending respectively on their chemical structure and presence of crosslinking, the medium temperature and Tg and Tm values of the polymer, the presence or absence of load, and the rate at which it is applied. As a result of this the polymer could function differently than intended, and one needs to know these property changes before converting the polymer into a product. Rubber elasticity When a solid polymeric compound softens upon increase of medium temperature, it still retains the interactions between the chains of a coil and between coils. Upon increase of temperature the chains in a coil and the coils themselves become more and more mobile with respect to each other and few

Control of polymeric biomaterial surfaces

35

contacts are left intra- and intermolecularly. If the temperature is not high enough to completely melt the polymer or if there are covalent linkages (crosslinks) between the chains, it first responds by deforming and then by regaining the original form when subjected to a short deforming force. A change in the position of coils with respect to each other does not take place. This behavior of the material, represented by a strong deformation under a small deforming force and elastic recovery was first observed with natural rubber and is known as rubber elasticity. Relaxation processes When the polymer is brought above its softening point and if the application of the deforming force is sufficiently long, the chains in the structure have time to reorganize with respect to each other into an energetically more favorable state. In this way, the inner tension that has developed in the system is relieved by the process called `relaxation'. If the viscosity of the system is low, the relaxation is too rapid and no recovery can take place. If the viscosity is high the relaxation time is longer and there is a chance for recovery. The latter type of materials flow if the deformation is prolonged, but they are elastic if the deformation is short and rapid. A material that shows both rubber elastic behavior and relaxation is called viscoelastic.

2.4

Polymer-solvent interactions

Polymers present certain chemical and physical properties when dry, but upon coming in contact with liquids all this changes. Certain regions or phases of the polymer might interact with the liquid more than the rest of the structure leading to a different surface composition than in a dry state. The liquid could also lead to swelling or dissolution of the polymer, changing the properties further. It is therefore very important to understand the behavior of the polymer in a variety of solvents.

2.4.1 Polymer gels, hydrogels Crosslinked polymers by definition are composed of a large number of chains strongly (mostly covalently) bonded to each other. Upon immersion of the crosslinked polymer in a solvent the chains forming the structure separate from each other and become solvated but cannot dissolve away due to the restrictions imposed by the links between the chains. This structure is called a gel. If the solvent in which the structure is introduced is water, then the water-swollen structure is called a hydrogel. There are cases where the polymer is introduced to a `less good' solvent, a solvent in which the chemical structures of the solvent and the polymer are not

36

Surfaces and interfaces for biomaterials

similar, and solvation is not as easy as it is in a good solvent. Under those circumstances, an uncrosslinked polymer swells but it is not possible to separate the chains from each other due to the weak interaction with the solvent. In these cases one still obtains gels even though there are no crosslinks. In the latter case, it is possible to achieve solubility by increasing temperature, but a truly covalently crosslinked system cannot dissolve upon increase of temperature.

2.4.2 Sol-gel transition The sol phase is defined as a flowing fluid, whereas the gel phase is non-flowing and maintains its integrity against external forces. Above the critical concentration (critical gel concentration, CGC) of a polymer, the gel phase appears. The CGC is most often inversely related to the molecular weight of the polymer employed. The development of physical junctions is needed for gelation, which must be sufficiently strong to overcome the dissolving forces of the solvent. Thermoreversible gelation of gelatin and polysaccharides such as agarose, amylose, and amylopectin, cellulose derivatives, carrageenans are well known. At high temperatures, they are assumed to have a random coil conformation. On reducing the temperature, they start to form double helices and aggregates. Most natural polymers form a gel phase on lowering temperature. They are said to have Upper Critical Solution Temperature, UCST. However, aqueous solutions of some cellulose derivatives exhibit reverse thermogelation (gelation upon increase of temperature) and are said to have a Lower Critical Solution Temperature, LCST. Cellulose itself is not soluble in water, however, it becomes water soluble upon introduction of hydrophilic moieties (Jeong et al., 2002). When these derivatives have a balance between hydrophilic and hydrophobic moieties, they undergo sol-to-gel transitions upon change of temperature. The temperature at which sol-gel transition takes place depends on the level and location of the substitution. Upon increasing temperature, water becomes a poorer solvent and polymer-polymer interactions become dominant and a gel forms. Table 2.1 shows the LCST values of some polymers.

2.4.3 Influence of bulk and surface properties Creation of polymers with different bulk and surface properties can be achieved in a variety of ways for applications within a wide range of industries. For polymeric materials, the surface energy value is determined mainly by the chemical structure at the surface. It has been suggested that amorphous, comb like polymers possessing a flexible linear backbone onto which are attached side chains with intermolecular interactions exhibit low surface energy values. Surface energy is influenced by parameters such as roughness, the nature of the

Control of polymeric biomaterial surfaces

37

Table 2.1 Polymers with a LCST in water (adapted from Jeong et al., 2002) Polymer

LCST (ëC)

Poly(N-isopropylacrylamide), NIPAM Poly(ethylene glycol), PEG Poly(propylene glycol), PPG Poly(methacrylic acid), PMAA Poly(vinyl alcohol), PVA Poly(vinyl pyrrolidone), PVP Methylcellulose, MC Hydroxypropylcellulose, HPC Poly(N-vinylcaprolactam)

32 120 50 75 125 160 80 55 30

polymer backbone and the pendant chain. For example, grafted perfluorocarbons with varying lengths influence surface properties in a very significant manner (Barbu et al., 2002).

2.4.4 Influence of composition (monomers, polymer type, additives) Polymer-solvent interactions depend on the chemical properties of the solvent, the ingredients of the polymer and the properties of the medium. Every system tends to reduce its internal energy (
U) or enthalpy (
H) and increase its entropy (
S). In other words every system tends to decrease its Gibbs Free Energy (
G). The relation between free energy, enthalpy and entropy at a certain temperature is given as:
G ˆ
H ÿ T
S When
G is decreased a process is spontaneous. During the dissolution process there is an increased mobility of the solute molecules and thus there is an increase in entropy. Enthalpy is also expected to decrease because there should be increased interaction between the solute and the solvent molecules if there is a good solvent for the solute. On the whole, free energy is decreased (
G < 0), making the dissolution process spontaneous. The enthalpy change is an outcome of the solvent-solute interaction. When there are more and stronger interactions leading to decreased enthalpy, this leads to increased solubility. The major factor at this stage is the chemistry of the monomer chains and of the resultant polymer. For example, if the polymer is composed of hydrophilic monomers such as N-vinylpyrrolidone, than the interaction with water will be strong and the enthalpy will decrease, leading to decreased free energy and increased solubility. If, however, the molecule in question is styrene, then the polymerwater interaction will be less than polymer-polymer and water-water interactions. Thus, the result would be insolubility. Only upon increase of

38

Surfaces and interfaces for biomaterials

temperature is mobility of chains increased leading to higher entropy and thus eventually a temperature is reached where solubility is achieved.

2.5

The polymeric surface and surface-bulk difference

Polymers are synthesized to yield glassy solids or powders. It is not easy to distinguish between the surface and the bulk. Upon conversion of the polymer into a product, the bulk and surface become different both in terms of chemistry and topography. There are a variety of reasons for the bulk and surface chemistry to differ: · oxidation of the surface · orientation of the macromolecules in a way that gives better interaction with a mould · different free energies of the various components of the macromolecule. Upon introduction into a liquid medium, the differences become more distinct. The interaction of the polymer chains with the solvent modifies the surface and bulk in terms of concentration and localization of hydophilic and hydrophobic groups. Upon introduction to a hydrophilic liquid, the concentration of hydrophilic groups or segments on the surface increases while in the interior hydrophobic groups concentrate. As a result, the surface properties of the polymer become different from those of the bulk. The packing densities on the surface and in the bulk are different. Due to thermodynamic constraints, the hydrophobic groups and segments form more hydrophobic interactions that lead to an increase in the packing density while on the surface, interaction between the solvent and polymer overcomes the polymer-polymer attraction leading to a swollen or highly extended conformation, and a lower packing density. In good solvents, the polymer-solvent interaction is more favorable than the polymerpolymer one. Polymer chains in good solvents swell due to steric repulsion. This spatial size of a polymer coil is much smaller than its extended contour length but larger than the size of a typical chain. The reason for this peculiar behavior is entropy combined with the favorable interaction between polymers and solvent molecules in good solvents. Similarly, for the adsorption of polymer chains on solid substrates, the conformational degrees of freedom of polymer coils lead to salient differences between the adsorption of polymers and small molecules. In the case of `poor' solvent conditions, the effective interaction between polymers is attractive, leading to collapse of the chains and to their precipitation from solution (phase separation between the polymer and the solvent). In this case, the polymer size decreases, like any space filling object embedded in threedimensional space. Thus if the polymer is a homopolymer with no segments of different functionalities, then in the poor solvent the packing on the surface is expected to increase in comparison to the chains in the inside.

Control of polymeric biomaterial surfaces

2.6

39

The general properties of a biomaterial surface

In the past, biomaterial biocompatibility was considered to be passive behavior towards the biological system. This requires the biomaterial to be nonthrombogenic, non-allergenic, non-carcinogenic, and non-toxic (Klee and HoÈcker, 2000). Williams defined biocompatibility, however, as `the ability of a material to perform with an appropriate host response in a specific application' (Williams, 1999). A variety of properties determine biocompatibility. Among these are the mechanical and chemical/physical properties of a material, both in the bulk and at the surface; the surface exerts the heavier influence because it is the component where the biomaterial and the biological system meet and interact. The properties of the surface that have utmost importance are: · the chemical structure ± hydrophilicity ± presence of groups that could initiate reactions in the biological system · the morphology ± the distribution and abundance of hydrophilic/hydrophobic and crystalline/amorphous phases ± surface topography, i.e. surface roughness and the presence of physical forms. The surface characteristics of a polymer are considerably different from the bulk characteristics. In a dry state, due to the minimization of surface energy and chain mobility, the non-polar groups move to the phase boundary with air while the reverse happens in an aqueous medium. Following this, low molecular weight components could migrate either towards or away from the surface leading to differences in the properties of the surface and the bulk. At the phase boundary between the biomaterial and the aqueous surrounding tissue, a different situation arises than at the phase boundary between the biomaterial and air. When the implant comes into contact with the biological system the following reactions are observed: 1.

2.

Within the first few seconds molecules (especially proteins and lipids) from the surrounding body liquids and tissue are deposited. In blood the properties of this adsorbed protein layer determines hemocompatibility, while in tissues it controls further reactions of the cell system. The nature of the adsorbed proteins is dependent on the surface characteristics of the implanted material. Upon implantation, the tissue neighboring the implant undergoes mechanical, chemical or physical damage or a combination of these. This damage could be acute (short-term and intense) or chronic (long-term) depending on the degree of stability of the material surface. In any case a period of inflammation is observed.

40 3. 4.

5.

Surfaces and interfaces for biomaterials A biocompatible implant, if inert, is surrounded by a thin layer, mainly consisting of collagen. A biocompatible implant, if not stable will undergo some sort of degradation and/or erosion in the harsh biological medium. This will continue until the implant is completely removed from the medium. A nonbiocompatible implant, if not stable and inert, will undergo degradation and erosion but its products will initiate a variety of undesirable reactions.

Since the surface is the outermost region of the implant interfacing tissue, it should fulfill the requirements of biocompatibility.

2.7

Modification of polymer surfaces

Surfaces of polymers are modified intentionally and unintentionally by various mechanisms and approaches as described below.

2.7.1 Oxidation by air Oxidative degradation is a route for unintentional surface modification and plays an important role in the aging of polymers. Since mainly the surface of a polymer is exposed to air, the effect of oxygen is observed more on the surface than the bulk. This oxidation is influenced by light, water, moisture and temperature. With most polymers, the exterior signs of oxidation are yellowing and if oxygen penetration was possible, then as an increase in brittleness. Unsaturated polyolefins are susceptible to oxidation by air. Oxidation generally induces the formation of peroxide which later leads to chain scission and probably simultaneous free radical formation. This decrease in chain length causes a deterioration of mechanical properties. Thus upon exposure to air chain length decreases and surface oxidation leads to more polar groups than initially. The extent of this change in the bulk is very limited and is controlled by the gas permeability of the solid polymer.

2.7.2 Grafting Grafting can be achieved in a variety of ways leading to products with different forms and chemistries, such as interpenetrating, crosslinked, brush, stimuli responsive, etc. Interpenetrating graft Interpenetrating graft copolymers are obtained where a polymer is dissolved in the monomer to be grafted and then allowing polymerization to take place. The

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chains of the earlier polymer will be entrapped in the newly formed structure. If this new polymer has functional groups or more unsaturations than one, the newly formed structure will be crosslinked within itself more strictly entrapping the original polymer. When surface modification via crosslinked grafts is considered, then the approach would involve wetting or swelling of the surface of the polymer by the monomer and then allowing polymerization to take place. The thickness of the grafted region will vary with the extent of wetting of the surface before initiation of polymerization. In this manner the surface of a mechanically suitable hydrophobic polymer could be rendered hydrophilic for increased hemocompatibility or for decreased lipid adsorption. Crosslinked graft Crosslinked graft copolymers are obtained where the unsaturated polymer is dissolved in the monomer to be grafted and allowing polymerization to take place. The newly formed chains could terminate by a reaction with the unsaturations on the chains and lead to crosslinkages. When surface modification via crosslinked grafts is considered then the approach would be as in the case of the interpenetrating graft. Brush graft Brush grafting could be achieved by activating the functional groups on the polymer to react with newly introduced monomers leading to brushes on the surface. For example, in a typical application, hydrophilic polymer brushes bearing alcoholic hydroxyl groups were introduced onto a polyethylene (PE) surface by radiation-induced grafting of 2-hydroxyethyl methacrylate (HEMA), vinyl acetate (VAc), and glycidyl methacrylate (GMA). HEMA already carried a hydroxyl group. VAc or GMA grafted membranes, on the other hand, were hydrolyzed to yield the hydroxyl groups. Thus, a polyethylene surface was made hydrophilic by brush grafting. The brush presence also modified the properties of the material. When adsorption of gamma globulin onto this material was tested it was observed that upon increase of the amounts of the hydrophilic polymer brushes, adsorption of gamma globulin on the PE base was decreased. Also, these hydrophilic polymer brushes bearing multiple glycol groups (from GMA) and alcohol groups (from VAc) provided an attractive site for the linkage of ligands using chemical methods, a property which was not available for untreated PE (Kawai et al., 2003). Stimuli-responsive grafts Stimuli-responsive polymers respond to small changes in their environment with responses that can be used in various applications (Jeong and Gutowska, 2002).

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The most frequently applied changes or stimuli are temperature, pH, and ionic strength. The responses of polymers grafted on surfaces are presented when introduced into solutions. Recent advances in the design of stimuli-responsive polymers have created opportunities for novel biomedical applications. The responses can be observed as shape, surface characteristics, solubility and as enabled sol-gel transition. The extent of the response needs to be controlled, and therefore the reasons (mostly thermodynamic) for them have to be understood. A typical, and most commonly used, stimulus (temperature) responsive polymer is NIPAM (N-isopropyl acrylamide). NIPAM is soluble below 32 ëC and precipitates above 32 ëC in water. Below the phase transition temperature, i.e., LCST, the hydrogen bonds between the polymer and the water molecules are favored, and thus the polymer stays soluble. The chains are fully extended. Above this temperature the hydrogen bonds are ruptured due to excessive thermal motion of the molecules and the polymer chains prefer to interact with each other leading to insolubilization. The properties of this polymer could be modified by copolymerization with other monomers. Copolymerization of NIPAM with butylmethacrylate for example, decreases the LCST of aqueous copolymer solution because the co-monomer is hydrophobic. Copolymerization with hydrophilic co-monomers on the other hand results in an increase in LCST thus enabling modified responses of a system constructed of NIPAM. If two different stimuli-responsive hydrogels are brought together (i.e. NIPAM and Nvinyl caprolactam) in a hydrogel then two different response ranges (such as two different inversion temperatures) can be obtained. The changes in chain extension levels are also important in certain tissue engineering applications. When the temperature of the medium was decreased below 32 ëC, confluent cardiac myocyte layers were lifted from cell culture dishes without need for trypsinization, an enzymatic treatment which might be harmful to the resultant cell culture if not used properly. The sol-to-gel transition temperature of PEG±PLGA±PEG triblock copolymers was shown to be controlled over a temperature range of 15 ëC to 45 ëC in aqueous solution by changing PLGA and PEG length and the ratio of lactic to glycolic acid (Jeong and Gutowska, 2002). Owing to the triblock topology, PEG±PLGA±PEG polymers have limitations in terms of molecular weight and degradation profile and the ability to show the sol-to-gel transition in a desired range of ~10 ëC to 30 ëC. Graft copolymers of PEG±g±PLGA and PLGA±g±PEG, however, present sol-to-gel transitions at ~30 ëC. PEG±g±PLGA copolymers have hydrophilic backbones and form gels with short durability, whereas PLGA±g±PEG copolymers have hydrophobic backbones and form much more durable gels. Poly(L-lysine)±g±poly(histidine) is an example of a pH sensitive or responsive system (Jeong and Gutowska, 2002). Poly(L-lysine) is positively charged at physiological pH. The pKa of poly(histidine) is 6.0, and it therefore undergoes conformational changes at pH lower than this. Poly(propyl acrylic

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acid) (PPA) and poly(ethacrylic acid) (PEA) are also sensitive to pH changes in the environment. Both polymers have the capability to ionize at high pH and are neutral at lower pH. Thus, they are soluble at high pH and insoluble or less soluble at lower pH. This is reflected in the form of abrupt conformational changes at pH around 5 to 6. Certain polymers have the capability to hypercoil or to form hydrophobic bonds to create compact molecules, and this makes it possible to induce the macromolecules to change their conformation in response to local stimuli. Polymers with weakly charged pendant groups, i.e., either weak acids or bases, exist in the form of extended chains due to repulsion between these charged groups. If, in addition, the polymer also bears alkyl or aromatic pendant groups, then the latter will initiate hydrophobic interactions which will lead to the hydrophobic groups being located in the interior of the structure, thus enabling maximal hydrogen bonding between the polar groups and the water molecules. This `hydrophobic effect', is also the principal driving force in the formation of lipid-based assemblies such as cell mebranes and in determining the conformation of native proteins. Once medium properties are altered, the conformation of the molecule can also be reversibly changed (Tonge and Tighe, 2001).

2.7.3 High-energy treatments Polymer surfaces can be activated by application of high-energy radiation such as glow discharge plasma or gamma irradiation. Plasma modification Low-temperature plasmas are produced by electrical discharge in low-pressure gases. They consist of a mixture of highly reactive species, i.e., ions, radicals, electrons, photons and excited molecules. Their chemical composition and physical characteristics are determined, in addition to the gas used, by device parameters, such as chamber geometry, gas flow rate, frequency and the power applied. This method is used to modify the chemistry and morphology of polymer surfaces to a depth of several tens of microns thus leaving the bulk properties practically intact. Thus, completely different chemistry, hydrophilicity or surface roughness can be obtained. Riccardia et al. (2003) used air as the gas to achieve plasma treatment of polyethylene terephthalate. The surface chemical and physical modification on PET fibers was a remarkable increase in hydrophilicity, extensive etching and a low molecular weight. XPS studies revealed formation of C±O and C±N bonds in the surface layer and the simultaneous decrease of C±C and C±H bonds. They interpreted etching as mainly a consequence of ion bombardment, while surface chemical modifications were mainly due to the action of neutral species on the

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plasma-activated polymer surface. The main effects of the active species created from the gas on polymer chains is mainly chain scission with new functional group creation and crosslinking. Grafting of nitrogen-carrying functional groups by plasmas requires nitrogencontaining gases. This is achieved through generation of nitrogen-containing radicals or excited N2, atomic nitrogen, NH, NH2 radicals, etc. Grafting of amino groups by plasmas is generally known to occur in combination with that of other nitrogen functionalities (Meyer-Plath et al., 2003). In another study polyethylene (PE) surfaces were treated with corona discharge with power ranging from 10 to 50 W (treatment duration 5 s) using air as the gas; the wettability of the PE sheet was significantly increased as judged by the water contact angle. New C±O, C=O and O±C=O bonds were detected with XPS (Lee et al., 2003). Gamma irradiation One of the most common irradiation types in industrial use is gamma. This is mainly produced by a Cobalt-60 source. Their main advantage is that they are very penetrating. Crosslinking and chain scission are the two major effects that gamma radiation has on polymers. This exposure could be applied for the purpose of processing or for sterilization. Radiation grafting and hydrogel formation are among the more important gamma irradiation applications (Clough, 2001). Grafting is employed because with it the surface properties can be tailored according to needs, while the material retains its bulk properties. Some outcomes are an improvement of chemical resistance, wettability, biocompatibility and hemocompatibility, dyeability of fabrics and antistatic properties. They could also be made to attach functional groups to enable the material to immobilize enzymes and other bioactive species. The advantage of radiation crosslinking over chemical crosslinking is that the former is more economical, faster, allows substantial decrease in the chemicals in the material processed (thus lower risk of bioincompatibility or allergenic response), and can be carried out at around room temperature. When irradiated, polymers are either inclined to undergo crosslinking or chain scission depending on the chemistry and the processing conditions. A polymer that undergoes chain scission could crosslink when the conditions are modified. Alterations in surface morphology can occur upon irradiation, especially in the form of an increase in roughness, a change that alters properties such as contact angle and cell adhesion.

2.7.4 Self assembled monolayers (SAMs) Self assembled monolayers are spontaneously formed highly ordered structures guided by the surface on which attachment takes place, by the components of the macromolecule and of the solvent. A well known example is alkane thiols on

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gold surfaces where the thiols attach to the gold surface with alkane tails sticking out and eventually align to form a tightly packed crystalline structure. The properties of this assembled layer depend very much on the alkane length. PEI (polyethyleneimine)-PLGA block copolymers were also found to be selfassembled in water where the PLGA segment acted as a hydrophobic aggregate block and the PEI segment as a hydrophilic corona-forming block (Nam et al., 2003). The block copolymers formed micelle-like aggregates in water and the size of the aggregates depended on hydrophobic block length and the ionic state of the hydrophilic block. The aggregate size decreased when the PLGA block length decreased and the PEI block was protonated. The amphiphilic nature of block copolymers consisting of hydrophilic and hydrophobic blocks provides an opportunity to form micelle-like structures in water. However, they do not form if they are introduced directly to an aqueous medium. They form their micellelike aggregates only when the solution of the polymer in an organic solvent such as DMF is introduced into aqueous media and then subsequently dialyzed, allowing time for organization in the form of micelles.

2.7.5 Surface patterning The use of chemical and physical patterns on polymer surfaces is on the rise due to the variety of uses they can be put to. Among them are biomedical uses such as construction of biosensors and tissue engineering applications. Patterning can be achieved by various methods as described in the following paragraphs. Contact printing/microprocessing This is basically a photolithographic approach. A variety of processes such as embossing, fused deposition modelling, ink jet printing, microcontact printing, proton micromachining and rapid prototyping are among the methods that could be classified under this heading. Embossing The first step is to fabricate a master die using electron beam lithography from which all other devices will be made. This master is made using a quartz substrate which has been coated with a thin metal film layer of Ti/Pd/Au to remove charge during exposure and to act as a plating base (Casey et al., 1997). A negative e-beam resist which facilitates the writing of small features over a large area in a relatively short time is used. The resist is developed to produce the pattern which is intended to be transferred into the plastic. The sample is then cleaned using an organic solvent to remove any remnants of resist or primer. It is now electroplated with nickel to the required depth which is typically a half to a third the height of the resist. The resist is removed by

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refluxing in a solvent after which the sample is a nickel die ready to use as an embossing master. If the features required for embossing are less than 150 nm in size, then the master is made by dry etching the resist pattern into quartz using an intermediate titanium layer. The quartz is then etched. The polymer used for embossing could be cellulose acetate which is in l mm thick sheet form. The important characteristics of this plastic are that it has a smooth, gloss surface with good clarity and a relatively low flow temperature in the range 135±175 ëC. The embossing process requires application of a constant force to the die/plastic/ platen stack. It is essential that a constant pressure is maintained throughout not only at the imprinting stage of the procedure, but also during the cooling of samples. Standard embossing times are 30 minutes at the elevated temperature of 135 ëC, followed by a 15-minute cooling period. Fused deposition modelling (FDM) process The FDM method forms three-dimensional objects from computer-generated solid or surface models like a typical rapid prototyping process. Models can also be derived from computer tomography scans, magnetic resonance imaging scans or model data created from 3D object digitizing systems. FDM uses a small temperature-controlled extruder to force out a thermoplastic filament material and deposit the semi-molten polymer onto a platform in a layer by layer process. The monofilament is moved by two rollers and acts as a piston to drive the semimolten polymer. At the end of each finished layer the base platform is lowered and the next layer deposited. The designed object is fabricated as a threedimensional part based on the deposition of thin layers of the polymer. The deposition path and parameters for every layer are designated depending on the material used, the fabrication conditions, the applications of the designed part and the preferences of the designer. This method was employed in the production of novel scaffolds with a honeycomb-like pattern, fully interconnected channel network, and controllable porosity and channel size (Zeina et al., 2002). Poly(-caprolactone) (PCL) was developed as a filamentmodeling material to produce porous scaffolds, made of layers of directionally aligned microfilaments. The PCL scaffolds were produced with a range of channel sizes 160±700 m, filament diameter 260±370 m and porosity 48± 77%, and regular geometrical honeycomb pores, depending on the processing parameters. Ink jet printing technique The technique is very simple. The polymer is dissolved in a volatile solvent (e.g. chloroform, trichloroethylene) and the liquid is ejected by the printing head on the substrate (i.e. alumina). As the drop reaches the surface the solvent evaporates leaving the solidified polymer behind. If more than one drop is

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ejected the solidification time is increased, leading to wider, thicker and more homogeneous polymer deposition (Pede et al., 1996). Microcontact printing of DNA In this method a rubber stamp that carries the inverse of the required pattern is immersed in the patterning solution (DNA, protein, enzyme, antibody, etc.) and then applied on the surface of the target material. The pattern is generally produced by photolithography on silicone based rubbers. In a typical application, photoresist patterns of 5 m line and 10 m space were prepared by photolithography on a silicon substrate which was used as a master for the stamp (Fujita et al., 2002). Patterned poly(dimethylsiloxane) (PDMS) stamps were made by mixing silicone elastomer and a curing agent and pouring the mixture in a plastic disposable container with the silicon master at the bottom. This was then heated in an oven, and the hardened polymer removed. A dilute solution of 3-(2-aminoethyl amino)propyl triethoxy silane (amino silane) was applied directly to the stamp and was microcontact printed on mica substrates. DNA solution was applied on the patterned substrate, washed with pure water, and dried under nitrogen gas thus creating a pattern of DNA on the mica surface. Proton micromachining In this application a beam of protons is focused on the target and scanned to create the pattern needed. A resolution of 1 m could be achieved by this approach. Sanchez et al. (1999) used a piece of PMMA as the template material, exposed to a 0.6 MeV proton beam using a nuclear microscope. The beam spot was scanned over the PMMA in a series of patterns consisting of nine different types of ridge/groove structures. Multiple repeated exposures of the patterns were carried out to achieve a homogeneous exposure. Rapid prototyping This system involves fabrication of scaffolds for tissue engineering applications via robotic desktop rapid prototyping. In one recent application the experimental setup consisted of a computer-guided desktop robot and a pneumatic dispenser (Ang et al., 2002). As dispensing material, chitosan and chitosan/hydroxyapatite (HA) dissolved in acetic acid was chosen. It was forced out through a small nozzle into a dispensing medium which was sodium hydroxide and ethanol. Layer-by-layer, the chitosan was fabricated with a preprogrammed pattern. Neutralization of the chitosan formed a gel-like precipitate and the hydrostatic pressure in the sodium hydroxide (NaOH) solution kept the cuboid scaffold in shape. A good attachment between layers allowed the chitosan matrix to form a fully interconnected channel architecture.

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Laser etching (laser photoablation) Laser photoablation or laser etching is a highly efficient and versatile etching process by which selected regions in a polymer layer can be removed by highenergy UV laser irradiation. The material is removed in a one-step process, requiring no developing or other steps before or after exposure. In addition to a wide variety of polymeric materials, photoablation has also been successfully used for laser-etching of several metals, oxides and other inorganic materials. The largest applications of photoablation include drilling of vias in multilayer microelectronic circuits, fabrication of nozzles in inkjet printheads, and corneal shaping for vision surgery. Numerous other applications of ablation have been implemented and many others are being researched in a variety of electronic, optoelectronic and medical fields. In the laser ablation (LAB) process using high-energy ultraviolet laser beams, polymers are chemically decomposed and removed without heating, therefore, smooth bottoms and sidewalls of the etched holes can be obtained. In practice, a polymer film is glued on a Si substrate and is etched by the LAB process. Slant planes can be easily formed by the relative movement of the laser beam. In a typical system the light source is a KrF excimer laser operating at a wavelength of 248 nm. The laser beam radiating from the source goes through a mask and a pattern on the mask is imaged on a sample. The sample is etched vertically in the irradiated area and the depth of the etched hole is proportional to the number of laser pulses. Laser photoablation can be carried out either in a spot-by-spot fashion using a focused beam from a low-power laser and raster-scanning it to address all the locations, or by projection-imaging the desired pattern of vias, lines, etc., from a mask onto the substrate using a high-power laser, thereby ablating thousands of features simultaneously. In the case of near-IR (NIR) laser patterning, however, it is believed that laser etching mostly results in substantially disordered profiles due to thermal effects. It is accepted that the thermo-mechanical properties of polymer materials play an important role in determining the quality of NIR laser etched surfaces. E-beam etching Conventional lithography involves patterning with light, generally in the UV region, of a thin polymer film that is spin-coated onto a silicon wafer. The radiation crosslinks (negative resist) or degrades (positive resist) the exposed polymer which is then washed (developed) to remove or retain the design, leaving behind a patterned polymeric surface. The same approach could be applied to a polymer with a significant thickness to create a pattern on the surface while retaining the bulk properties. This could be achieved by e-beam irradiation. The e-beam machine plays a significant role in the processing of polymers in a similar fashion. A number of different designs and energies are available.

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Industrial e-beam accelerators with energies in the 150±300 keV range are used when low penetration is needed. Accelerators operating in the range 1.5 MeV are used where more penetration is needed. High-energy commercial e-beams (10 MeV) are used in the sterilization of medical supplies already boxed and ready for shipment. They have high dose rates and therefore short processing times. Low penetration indicates more effective use of e-beam energy than with gamma. In patterning 100 keV electron beams a magnetic lens is used. Patterning is achieved by using an electron scattering mask. E-beam is capable of patterning at around 70 m resolution (Clough, 2001).

2.7.6 Immobilization of functional groups and molecules Chemical modification of the surface through introduction of new groups to the polymeric product is designed to impart certain properties to a material. This does not necessitate the modification of the bulk. In a typical application, thiocyanation of plasticized PVC led to a product with increased hydrophilicity but also significantly reduced retention of two types of bacteria strongly implicated in implant related infections, S. aureus and S. epidemidis (Nirmala and Jayakrishnan, 2003). Immobilization of heparin onto the surface of biomaterials to render them hemocompatible has been used in a large number of studies. Heparin immobilization was reported with PVA, methyl methacrylate, and polyurethane to name a few. Heparin was immobilized on Gore-TexÕ to improve its vascular graft performance (Begovac et al., 2003), to introduce hemocompatibility to nonwoven fabrics of polypropylene (Tyan et al., 2002), to silicone rubber, polyethylene, polypropylene and polyvinylchloride (Michanetzis et al., 2003), to PLA (Zhu et al., 2002), etc. Another approach is surface molecular imprinting. This technique helps introduce specific recognition sites into polymers. It has the ability to be applied to a wide variety of target molecules. If a high interfacial activity functional molecule is used, the enzyme-mimic site can be located on the surface of the polymer. In a specific example, the technique was used to help resolve optically active amino acids (Toorisaka et al., 2003). 2-Methacryloyloxyethyl phosphorylcholine (MPC) polymers were synthesized to mimic the biological membrane structure. MPC polymers are useful for surface modification of conventional materials even when random copolymers composed of MPC and alkylmethacrylate are applied as coating polymers. They effectively reduce protein adsorption and denaturation and inhibit cell adhesion even when the polymer is in contact with whole blood in the absence of any anticoagulants. It was observed that phospholipid polymer surfaces showed excellent blood compatibility (Yamasaki et al., 2003).

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2.8

Surface analysis

A biomaterial surface presents different levels of crystallinity, roughness and chemical groups to the biological environment into which it is introduced. Immediately after implantation, the adsorption of a variety of biological molecules such as proteins and lipids change the surface further (Fig. 2.6). In order to be able to construct the ideal biomaterial, the surface has to be characterized and this is achieved by a variety of techniques. In order to determine the composition and structure of a biomaterial surface, different methods which provide varying degrees of information are commonly used. ATR-IR (attenuated total reflectance infra-red) or ATR-FTIR (attenuated total reflectance Fourier Transform infra-red) spectroscopy supplies the characteristic absorption bands of functional groups with an informational depth of 0.1±10 m. Samples with rough surfaces are studied with photoacoustic spectroscopy (PAS), which allows analysis down to approximately 20 m. The achieved informational depths are usually larger than the thickness of the modified interface which might have a several molecules thick modification as in-plasma modification approaches and therefore the spectra obtained could include peaks due to the bulk as well. X-ray photoelectron spectroscopy (XPS) (or ESCA, electron spectroscopy for chemical analysis) is a more surfacesensitive analytical method which supplies information not only about the type and amount of elements present but also about their oxidation state and chemical surroundings. Depths of approximately 10 nm can be achieved with this method (about 50 atomic layers). A more surface-sensitive method is secondary ion mass spectroscopy (SIMS) where primary ions interact with the polymer surface and the mass spectra of the formed ions (secondary ions) are obtained which give information about the chemical composition of the outermost atomic layers (approximately 1 nm in thickness). The application of atomic force microscopy (AFM) in comparison to scanning electron microscopy (SEM) delivers information about surface properties as far as molecular dimensions. Another advantage of AFM compared with SEM is that the sample is investigated in the original state (no sputtering).

2.6 A typical biomaterial surface in the in-vivo environment.

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The full characterization of the surface of a biomaterial is often verified only by application of several analytical methods to the sample. The depth of the various analytical systems could deliver information about are: SIMS and AFM about 1 nm, XPS about 10 nm and ATR-IR 4000 nm.

2.8.1 Morphology When modified in one of many ways, the morphology of a polymeric surface is also changed whether intended or not. In order to be able to study the extent and type of these changes, a variety of novel techniques are used. The following three microscopic techniques are among the most effectively and extensively used surface characterization techniques. Scanning tunnelling microscope (STM) Two researchers working at IBM in Switzerland invented in 1982 the scanning tunnelling microscope (STM), for which they were awarded the Nobel Prize for Physics. In this technique, a fine sharpened tip, which is presumed to have a single atom at the apex, is scanned above the surface of the sample. A voltage (bias potential) is applied between the tip and the sample and a small current (tunnelling current) which flows between the gap is measured. The variation in tunnelling current as it passes over the atomically corrugated surface is then recorded and, if this process is repeated across the entire sample, a threedimensional map of the surface can be obtained. The choice of materials investigated was partly limited by the requirement for them to be conducting, so that a tunnelling current could be measured. Atomic force microscopy (AFM) Atomic force microscopy (AFM) is a method increasingly being used to study and quantify surface properties of materials in their untreated, natural form. It is essentially a scanner that creates topographical maps of surfaces. A very sharp tip, located at the free end of a cantilever follows the contours of the surface as it is moved over it. The deflections of the tip are measured by an optical detector and recorded by a computer. This instrument is very similar to STM, except that the tip and the surface are in contact and interatomic van der Waals forces acting between them provide the contrast mechanism, rather than a tunnelling current. There was no longer a requirement for the sample to be conducting, thus opening the way for the study of many other materials (Smith et al., 1997). The data can be used to create a 2-D or 3-D image and measure surface roughness. It is very important in the biomedical sciences because most polymers are highly hydrated in vivo, and thus treatments for visualization as in SEM (high vacuum, conductive coats, etc.) would not maintain their actual surface

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topography. With AFM, however, it was possible to examine very fragile hydrogel materials like soft contact lenses in their swollen form and measure their surface roughness (Grobe et al., 1996). It is also possible to monitor real-time events like hydrolytic degradation. In a study by Davies et al. (1996) the degradation of the components of blends of poly(sebacic acid anhydride) (PSA) and PLA was studied revealing the preferential degradation of the PSA component leaving behind surfaces enriched in PLA. It was thus possible to expose PLA morphology. With the use of a similar technique, scanning force microscopy, close packed, needlelike organization of crystals of poly(butene-1) were shown with a resolution of the order of nanometers (Jandt et al., 1993). It was claimed that individual poly(butene-1) molecules could be observed. In another study, lamellar organization of polyethylene (Jandt et al., 1994) and various spherulitic surfaces (Lustiger et al., 1989; Harron et al., 1996) could be visualized. A similar observation was made by Hasirci et al. (2002) where degradation of PPF-NVP/ EGDMA reinforced PLGA was investigated revealing a lamellar organization of the bone reinforcement plates after a very brief incubation in distilled water. Scanning electron microscopy (SEM) SEM is a microscopic method extremely valuable in polymeric products and biomaterials areas along with many other research and application areas. Its main attributes are that it reveals the 3D topography of the specimen examined and its magnification could be extremely high (factors of hundred thousands). For polymeric materials, high magnifications are generally not possible because of the intensity of the electron beams damaging the thin polymeric samples which generally lead to deformation and even melting of the specimen. Other limitations of SEM are the need for a high vacuum which prevents visualization of solvated samples in their natural state and the need to coat the sample with conductive materials such as gold. In order to overcome these problems, methods of fixation or stabilization and obtaining durable replicas of the actual sample are tried with varying degrees of success. The resultant 3D image is generally worth the effort because it reveals a wealth of information about the specimen examined. SEM was used in a series of studies on drug delivery systems and bone plates with great success. A PLGA rod (or rather fiber with diameter ca. 0.65 mm) controlled pain relief system was designed to deliver analgesics for a duration of 3±4 weeks (Hasirci et al., 2003). The initial tests were carried out in rats by tying the fiber to the sciatic nerve before testing for the loss of sensitivity to external pain inflicting stimuli such as high intensity light. The fiber withstood implantation very well, without disintegrating with the only deformation being at the location where it was tied to the sciatic nerve. The PLLA-PPF based bone plates for fracture fixation in rats reveal that while one would perform its function of stabilizing a fracture properly the other implant would probably disintegrate by cracking under the applied stress (Hasirci et al., 2000).

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2.8.2 Chemistry The surface chemistry of a polymer intended for biomedical applications can be investigated by a number of approaches, mostly based on spectroscopy and surface energetics. In the following section, information about these methods is provided. X-ray photoelectron spectroscopy (XPS) ESCA is the abbreviation of electron spectroscopy for chemical analysis, and is the same as XPS (X-ray photoelectron spectroscopy). It is a method for studying the energy distribution of electrons ejected from a material that has been irradiated with a source of ionizing radiation such as X-rays. This powerful tool provides quantitative information about basic properties such as binding energy, charge, and valence state, which examines the atom as a part of its chemical environment. The impact of ESCA in polymeric characterization has been twofold: it can analyze relatively intractable materials without the need for special sample preparation and it is a surface sensitive method. In principle, when any material is bombarded by photons with energy greater than the binding energy of an electron in a given atomic shell or sub-shell, there is a finite probability that the incident photon will be absorbed by the atom and an electron is either prompted to move to an unoccupied level or ejected as a photoelectron. This depends on the energy of the incident photon and the atomic number of the target element. The kinetic energy of the photoelectron is: KE ˆ h ÿ BE where KE and BE are the kinetic energy and the binding energy of the photoelectron, and h is the energy of the incident photon. Although the X-ray may penetrate deeply into the sample to produce photoelectrons, most of these electrons lose energy in numerous inelastic collisions; only those atoms residing in the top few monolayers give rise to undistorted photoelectron spectra. The typical analysis depth in ESCA and in Ê , and they are truly surface analysis Auger electron spectroscopy is about 3±50 A methods. Auger electron spectroscopy (AES) In AES, an incident primary electron creates an excited ion near the surface which decays by the emission of a secondary Auger electron, whose kinetic energy is measured. As in XPS, the escaping Auger electron's kinetic energy limits the depth from which it can emerge, giving AES its high surface sensitivity and nanometer sampling depth. Auger images or maps can also be generated for specific elements with approximately 200 nm resolution. Auger

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finds its greatest strength in the analysis of inorganic materials not susceptible to electron-beam damage. Surface Raman Since its discovery in 1974, surface enhanced Raman scattering (SERS) has proved to be a very powerful and sensitive technique to investigate the vibrational properties of adsorbed molecules. It has been shown that the enhancement mechanisms can be of two kinds: a chemical phenomenon in which charge transfer between adsorbed molecules and the metallic substrate is involved and an electromagnetic enhancement which considers a localized surface plasmon. Yet, there is still some debate inside the SERS community as to whether one effect predominates or not (Grand et al., 2003). Since its original discovery, surface enhanced Raman scattering (SERS) has developed significantly towards detection of a single molecule. It was also more successful in biological applications than conventional Raman spectroscopy, especially in endoscopy and in-vivo diagnosis. Novel applications in biology include cancer gene detection, spectroscopy of living cells and single protein/ DNA detection. Progress in non-biological applications of SERS has been equally spectacular, with single molecule detection, spectroscopy of single dyes in large nanocrystals, or in carbon nanotubes (Etchegoin et al., 2003). Attenuated total reflectance (ATR) spectroscopy Before the advent of ESCA the only reliable method for polymer surface studies was infra-red spectroscopy (either attenuated total reflectance (ATR) or multiple internal reflectance (MIR) spectroscopy). This method requires fairly large samples with flat or easily deformable surfaces and typically gives information pertaining to surfaces and gives information penetrating to 1 m into the material. In recent years, Fourier transform infra-red (FTIR) has allowed analysis sensitivity to be improved, but even then FTIR can neither match the sensitivity of ESCA nor can it be as focused on the surface. In systems with variable composition within the range 0±1 m, the two methods could be applied to complement each other effectively. Infra-red microscopy Infra-red microscopy is a powerful technique that combines the image analysis capabilities of optical microscopy with the chemical analysis capabilities of infrared spectroscopy. The combination of these two techniques allows infra-red spectra to be obtained from microspectroscopic-sized samples. With the assistance of microscopy, samples as small as 0.01 g or even less (depending on the infra-red absorption characteristics of the components of interest) can be easily located and

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detected. Among its key roles are point identification or point mapping, using single element detectors to fingerprint defects and contaminants, analyze laminate structures, and to map chemical and physical property differences and gradients. These are critical to understanding product behavior, since chemical structure and conformation and physical property anisotropy at the microscopic level significantly influence the macroscopic performance of polymer products. A major limitation of FTIR microscopy has been the time needed using single element detectors to construct high spectral contrast maps of high spatial (lateral) resolution (~10 m) from large areas (e.g. >50 m  50 m). However, over the last few years, images based on mid-infra-red spectral differences and changes generated using FTIR microscopy instrumentation fitted with focal plane array (FPA) detectors have increasingly become available (Chalmers et al., 2002). In one application time-resolved FTIR measurements during isothermal crystallization of samples were carried out on a cast film. The crystallization of samples was traced by the C=O stretching band at 1722 cmÿ1 and the C±D stretching band at 2230 cmÿ1. Because of the large difference in molar absorbance coefficients, the C=O and C±D stretching bands were observed by separate measurements (He et al., 2002).

2.8.3 Energetics Surface energetics is very important in defining the biocompatibility of a polymer because this determines the type of molecules that do and do not adsorb on the surface. Tissue ingrowth, blood coagulation, cell damage to blood elements and immune signalling are all influenced by the energetics of the surface. Zeta potential A charged molecule in motion produces an electric field. The zeta potential is the electric potential of a charged particle at the plane of shear. The shear plane, or the plane of slip, is the distance from the surface to the distance in solution where the solvent molecules are not bound to the surface and are not moving as a unit with the particle. At this boundary, zeta potential can be determined. The surface potential is a very difficult parameter to characterize and it is far easier to determine zeta potential. It is measured at a shear plane near the particle surface and is therefore proportional to surface potential. It is affected by ionic strength; at high ionic strength, potential decreases much more sharply over the distance from the surface to the shear plane. This means a smaller zeta potential. It is therefore important to maintain a relatively constant ionic strength while characterizing the zeta potential of a dispersion as a function of pH. In the biomedical field, a large variety of applications are found. Gene therapy is one of them (Putnam et al., 2003). Complexation of plasmid DNA

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with polycations is a popular method by which to transfer therapeutic nucleic acid sequences to cells. One disadvantage of the approach is that the positive zeta potential of the complexes facilitates interaction with blood constituents, leading to serum protein adsorption and immune response. To circumvent this issue, investigators have developed polycations combined with polyethylene glycol (PEG) to create complexes with reduced protein adsorption potential. Contact angle The investigation of surface wettability by means of contact angle determination is of special interest in the characterization of the polymer surface. Contact angle may be geometrically defined as the angle formed by the intersection of two planes at a tangent to the liquid and solid surface at the perimeter of contact between the two phases and the third surrounding phase. Typically, the third phase will be air or vapor, although systems in which it is a second liquid essentially immiscible with the first are of great practical importance. If one considers the three-phase system where the liquid is designated as l, the surrounding gaseous medium as g, and the solid surface as s, then at equilibrium the contact angle  will be given by Young's equation as: gl
cos  ˆ sg ÿ sl where gl , sl and sg are the interfacial tensions at the respective interfaces. Contact angle measurements are carried out in various ways of widely differing sensitivity. Typically they are made with a goniometer and a syringe with a flat-tipped needle that is used to apply the solvent (generally double or triple distilled water) droplet on the surface. Advancing contact angles are recorded while fluid is added to the drop already on the surface. The values are generally reported as an average of five to ten measurements made on different areas of the sample surface. In culturing cells on biomaterial surfaces, contact angle is an important parameter that guides surface property modification attempts. For example, when oxygen plasma treated polyethylene surfaces were tested for tissue engineering purposes, the adherence of PC-12 cells were found to be higher as the water contact angle was lowered (higher surface wettability). However, this adhesion started to decrease as the contact angle was further decreased. Thus a bell-shaped curve for cell adhesion was obtained with the optimum water contact angle of 55 degrees (Lee et al., 2003).

2.9

Surface properties and biomaterials applications

Biomaterials are designed to augment, support or completely take over the functions of natural tissue or organs of human beings and are, therefore, designed to be biocompatible (non-toxic, non-immunogenic, non-inflammatory,

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non-carcinogenic) and to carry properties that match those of the tissue they are replacing. When functioning in the body the tissue-material interface is critical because this is where all the initial interactions such as contact with cells, proteins, and other biological entities takes place. The bulk of the biomaterial needs to satisfy the mechanical and physical requirements of the intended application, while an inert surface has to prevent immune responses, irritation or clotting. Biomaterial surfaces are, therefore, generally treated to improve their in-vivo responses. A typical example is the pyrolitic carbon-coated heart valve where the valve made of a metal or graphite is rendered hemocompatible and impact resistant by coating with a carbon layer. Heparin treated polymeric blood vessels and silicone sheathed sensor leads are all in this category. To improve blood compatibility it is a common approach to increase the hydrophilicity of the surface of the polymeric biomaterial and to achieve this surface modification by use of plasma treatment and subsequent grafting with a hydrophilic polymer. The optimum properties of a biomaterial can still not be elucidated, and a rule of thumb does not exist. The design of the surface of a biomaterial remains by far the most important and difficult task that a biomaterial scientist has to tackle.

2.10 Conclusion Surfaces of materials used in the biomedical field have always been important due to their proximity to the tissues. This is the first plane of contact between them and the nature of this contact is the main factor in determining their biocompatibility. Among the hottest current research trends today is nanobiotechnology where physical and chemical modification of biomaterial surfaces, their characterization and understanding of the nature of materialtissue interactions are of prime importance in today's biomaterials sciences. As the technology and our understanding of these increase, so will our ability to create better biomaterials. We will thus achieve higher quality of life.

2.11 References Ang TH, Sultana FSA, Hutmacher DW, Wong YS, Fuh JYH, Mob XM, Loh HT, Burdet E, Teoh SH (2002), `Fabrication of 3D chitosan±hydroxyapatite scaffolds using a robotic dispensing system', Materials Science and Engineering C, 20, 35±42. Barbu E, Pullin RA, Graham P, Eaton P, Ewen RJ, Smart JD, Nevell TG, Tsibouklis J (2002), `Poly(di-1H,1H,2H,2H-perfluoroalkyltaconate) films: surface organization phenomena, surface energy determinations and force of adhesion measurements', Polymer, 43, 1727±1734. Begovac PC, Thomson RC, Fisher JL, Hughson A, Gallhagen A (2003), `Improvements in GORE-TEX vascular graft performance by Carmeda BioActive surface heparin immobilization', Eur J Vasc Endovasc Surg, 25(5), 432±437. Casey BG, Monaghan W, Wilkinson CDW (1997), `Embossing of Nanoscale Features and Environments', Microelectronic Engineering, 35, 393±396.

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Chalmers JM, Everall NJ, Schaeberle MD, Levin IW, Lewis EN, Kidder LH, Wilson J, Crocombeg R (2002), `FT-IR imaging of polymers: an industrial appraisal', Vibrational Spectroscopy, 30(1), 43±52. Clough RL (2001), `High-energy radiation and polymers: A review of commercial processes and emerging applications', Nuclear Instruments and Methods in Physics Research, B, 185, 8±33. Davies MC, Shakesheff KM, Shard AG, Domb A, Roberts JC, Tendler SJB, Williams PM (1996), `Surface analysis of biodegradable polymer blends of poly(sebacicacid anhydride) and poly(DL-lactic acid)', Macromolecules, 29, 2205±2212. Etchegoin P, Maher R.C, Cohen L.F, Hartigan H, Brown RJC, Milton MJT, Gallop JC (2003), `New limits in ultrasensitive trace detection by surface enhanced Raman scattering (SERS)', Chemical Physics Letters, 375(1±2), 84±90. Fujita M, Mizutanic W, Gadd M, Shigekawab H, Tokumotoc H (2002), `Patterning DNA on mm scale on mica', Ultramicroscopy, 91, 281±285. Grand J, Kostcheev S, Bijeon J-L, de la Chapelle ML, Adam P-M, Rumyantseva A, LeÂrondel G, Royer P (2003), `Optimization of SERS-active substrates for near-field Raman spectroscopy', Synthetic Metals, 139(3), 621±624. Grobe GL, Valint PL, Ammon DM (1996), `Surface chemical structure for soft contact lenses as a function of polymer processing', J. Biomedical Materials Research, 32, 45±54. Harron HR, Pritchard RG, Cope BC, Goddard DT (1996), `An atomic force microscope (AFM) and tapping mode AFM study of the solvent induced crystallization of polycarbonate thin films', J. Polymer Science (Pt B): Polymer Physics, 34, 173±180. Hasirci V, Lewandrowski K, Bondre SP, Gresser JD, Trantolo DJ, Wise DL (2000), `High strength bioresorbable bone plates: preparation, mechanical properties and in vitro analysis', Bio-medical Materials and Engineering, 10(1), 19±29. Hasirci V, Litman AE, Trantolo DJ, Gresser JD, Wise DL, Margolis HC (2002), `Investigation of the interpenetrating network structure of a molecularly reinforced biodegradable implant', J. Materials Science, Materials in Medicine, 13, 159±167. Hasirci V, Bonney I, Goudas LC, Shuster L, Carr DB, Wise DL (2003), `Antihyperalgesic effect of simultaneously released hydromorphone and bupivacaine from polymer fibers in the rat chronic constriction injury model', Life Sciences, 73(26), 3323±3337. He W, Shanks R, Amarasinghe G (2002), `Analysis of additives in polymers by thin-layer chromatography coupled with Fourier transform-infrared microscopy', Vibrational Spectroscopy, 30(2),147±156. Jandt KD, McMaster TJ, Miles MJ, Petermann J (1993), `Scanning force microscopy of melt crystallized, metal evaporated poly(butene-1) ultrathin films', Macromolecules, 26, 6552±6556. Jandt KD, Buhk M, Miles MJ, Petermann J (1994), `Shish-kebab crystals in polyethylene investigated by scanning force microscopy', Polymer, 35(11), 2458±2462. Jeong B, Gutowska A (2002), `Lessons from nature: stimuli responsive polymers and their biomedical applications', Trends in Biotechnology, 20(7), 305±311. Jeong B, Kim, SW, Bae YH (2002), `Thermosensitive sol±gel reversible hydrogels', Advanced Drug Delivery Reviews, 54, 37±51. Kawai,T, Saito K, Lee W (2003), `Protein binding to polymer brush, based on ionexchange, hydrophobic, and affinity interactions', Journal of Chromatography B, Analyt Technol Biomed Life Sci, 790(1±2), 131±142. Klee D, HoÈcker H (2000), `Polymers for Biomedical Applications: Improvement of the

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Interface Compatibility', in Advances in Polymer Science, Springer-Verlag, Heidelberg, Vol. 149, 1±57. Lee SJ, Khang G, Lee YM, Lee HB (2003), `The effect of surface wettability on induction and growth of neurites from the PC-12 cell on a polymer surface', Journal of Colloid and Interface Science, 259, 228±235. Lustiger A, Lotz B, Duff TS (1989), `The morphology of the spherulitic surface in polyethylene', J. Polymer Science (Pt B): Polymer Physics, 27, 561±579. Meyer-Plath AA, Schroder K, Finke B, Ohl A (2003), `Current trends in biomaterial surface functionalization±nitrogen-containing plasma assisted processes with enhanced selectivity', Vacuum, 71, 391±406. Michanetzis GP, Katsala N, Missirlis YF (2003), `Comparison of haemocompatibility improvement of four polymeric biomaterials by two heparinization techniques', Biomaterials, 24(4), 677±688. Nam YS, Kang HS, Park JY, Park TG, Han S-H, Changa I-S (2003), `New micelle-like polymer aggregates made from PEI±PLGA diblock copolymers: micellar characteristics and cellular uptake', Biomaterials, 24(12), 2053±2059. Nirmala RJ, Jayakrishnan A (2003), `Surface thiocyanation of plasticized poly(vinyl chloride) and its effect on bacterial adhesion', Biomaterials, 24, 2205±2212. Pede D, Serra G, De Rossi D (1996), `Microfabrication of conducting polymer devices by ink-jet stereolithography', Materials Science and Engineering, C5, 289±291. Putnam D, Zelikin AN, Izumrudov VA, Langer R (2003), `Polyhistidine ± PEG:DNA nanocomposites for gene delivery', Biomaterials, 24(24), 4425±4433. Riccardia C, Barnia R, Sellib E, Mazzoneb G, Massafrac MR, Marcandallic B, Polettid G (2003), `Surface modification of poly(ethylene terephthalate) fibers induced by radio frequency air plasma treatment', Applied Surface Science, 211, 386±397. Sanchez JL, Guy G, van Kan JA, Osipowicz T, Watt F (1999), `Proton micromachining of substrate scaffolds for cellular and tissue engineering', Nuclear Instruments and Methods in Physics Research B, 158 185±189. Smith JR, Campbell SA, Mills GA (1997), `Probing atoms', Educ. Chem, 34(4), 107±111. Stachurski ZH (2003), `Strength and deformation of rigid polymers: structure and topology in amorphous polymers', Polymer, 44, 6059±6066. Tonge SR, Tighe BJ (2001), `Responsive hydrophobically associating polymers: a review of structure and properties', Advanced Drug Delivery Reviews, 53, 109±122. Toorisaka E, Uezua K, Goto M, Furusaki S (2003), `A molecularly imprinted polymer that shows enzymatic activity', Biochemical Engineering Journal, 14, 85±91. Tyan YC, Liao JD, Wu YT, Klauser R (2002), `Anticoagulant activity of immobilized heparin on the polypropylene nonwoven fabric surface depending upon the pH of processing environment', J Biomater Appl. 17(2), 153±178. Williams DF, The Williams Dictionary of Biomaterials, Liverpool University Press, 1999. Yamasaki A, Imamura Y, Kurita K, Iwasaki Y, Nakabayashi N, Ishihara K (2003), `Surface mobility of polymers having phosphorylcholine groups connected with various bridging units and their protein adsorption-resistance properties', Colloids and Surfaces B: Biointerfaces, 28, 53±62. Zeina I, Hutmacher DW, Tanc KC, Teoh SH (2002), `Fused deposition modeling of novel scaffold architectures for tissue engineering applications', Biomaterials, 23, 1169±1185. Zhu A, Zhang M, Wu J, Shen J (2002), `Covalent immobilization of chitosan/heparin complex with a photosensitive hetero-bifunctional crosslinking reagent on PLA surface', Biomaterials, 23(23), 4657±4665.

3 Organic thin film architectures: fabrication and properties M C P E T T Y , University of Durham, UK

3.1

Introduction

Molecular electronics is a fast moving and interdisciplinary subject that exploits the electronic and optoelectronic properties of organic and biological materials (Petty et al., 1995; Richardson, 2000; Tour, 2003; Maruccio et al., 2004). The areas of application and potential application are varied, ranging from chemical and biochemical sensors to plastic light-emitting displays. Molecular electronics also offers considerable scope for molecular nanotechnology, e.g., the development of electronic switching and memory structures operating, and addressable, at the molecular level. In most of these examples, the organic materials are required in the form of thin films (1 nm to 10 m). This presents a considerable challenge to materials scientists, as organic compounds, in their bulk form, can be fragile and difficult to handle. In this chapter, an overview of the more popular methods that may be used to fabricate thin layers of organic compounds will first be described. A distinction will be drawn between established deposition technologies, in many cases developed for use with inorganic materials, and those methods that actually allow molecular-scale architectures to be built-up on solid supports. Each deposition method differs in complexity and may be more suited to provide films in a particular thickness range. Specific types of organic compound are necessary for certain processes, e.g., self-assembly exploits the attraction between certain chemical groups. Some methods are `wet' (spinning) while others are inherently `dry' (plasma deposition). There are also implications for the degree of order and contamination levels in the deposited film. Following the introduction to organic film technologies, a brief review of some of the more powerful analytical techniques that can be used to reveal the structure and degree of molecular organization in the thin layers will be presented.

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3.2

61

Established deposition methods

3.2.1 Physical vapour deposition Solid materials vaporize when heated to sufficiently high temperatures, this process may proceed through the liquid phase. A thin film is then obtained by the condensation of the vapour onto a colder substrate (Maissel and Glang, 1970). This method has been used extensively to deposit films of inorganic materials, such as metals and their alloys. However, the technique is now being used for the formation of layers of low molecular weight organic compounds. Because of collisions with ambient gas atoms, a fraction of the vapour atoms will be scattered. For a straight line path between the evaporating material (source) and the substrate, it is necessary to use low pressures (< 10ÿ4 mbar) where the mean free path of the gas atoms is much greater than the sourcesubstrate distance. This allows the use of a shadow mask immediately in front of the substrate to define patterns. The low pressure also prevents contamination of the source material (e.g. by oxidation). Figure 3.1 shows a schematic diagram of a typical evaporation system. The system chamber, which can be made from glass or metal, is evacuated to a pressure of 10ÿ4±10ÿ6 mbar, normally with two types of vacuum pump, a rotary and diffusion pump, operating in series. The first step in physical vapour deposition requires the transformation of the condensed phase, solid or liquid, into the gaseous state. This conversion of thermal to mechanical energy can be achieved by a variety of methods. Resistive heating has been used to deposit fluorescent dyes, charge-transfer salts and large

3.1 Schematic diagram of a typical vacuum evaporation system for physical vapour deposition.

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macromolecules, such as the phthalocyanines (Petty, 2000; Park et al., 2004). Typical evaporation rates are 1±10 nm minÿ1. Other techniques include arc evaporation, RF heating or heating by electron bombardment. Deposition of polymer films by laser ablation is an area that offers some promise (Chrisey and Hubler, 1994). Laser pulsed methods have been used successfully for polyethylene, polycarbonate, polyimide, polymethylmethacrylate and, more recently, for the deposition of compounds for organic light-emitting displays and transistors (Salih et al., 1997; Hong et al., 2002; Blanchet et al., 2003). Materials that dissociate in the vapour phase may provide solid films with a stoichiometry that differs from that of the source. Therefore special techniques have been devised. One approach is to use the method of `flash' evaporation. It is also possible to evaporate from two, or more, sources and to control the flux from each to obtain a vapour with the required composition. This has been used effectively to deposit thin films of doped organic charge-transfer salts (Breen et al., 1993): one source is the charge-transfer salt, e.g., tetrathiafulvalene (TTF), while the other is the dopant, e.g., iodine. Laser co-ablation techniques can be used to deposit films of metal-polymer composites (Chrisey and Hubler, 1994). Molecular beam epitaxy (MBE) is a similar, but more expensive, variation of vacuum evaporation (Hara and Sasabe, 1995). However, an ultra-high vacuum (90% is rejected. In cell membranes, transport across the membrane occurs as a consequence of passive transport driven by the kinetic energy of the molecules being transported or by membrane transporters. Thus a molecule or ion that crosses the membrane by moving down a concentration or electrochemical gradient and without expenditure of metabolic energy is transported passively (diffusion). For charged molecules, the electrical potential across the membrane also becomes critically important. Together, gradients in concentration and electric potential across the cell membrane govern passive transport mechanisms. In cell membranes, active transport also takes place, this being the movement of molecular against an electrochemical gradient which requires energy expenditure. In membranes used for membrane separation processes, the volume flux produced may be expressed from consideration of the hydraulic resistance of the membrane used and the pressure drop over the membrane. In general the inverse of the resistance, hydraulic permeability being used. The permeability is defined by the pore size and structure and thickness of the membrane. Membranes can have different structures, they may consist of different, more or less cylindrical pores or may be composed of a packed bed of various small particles (which are of the order of nanometres). If the membrane consists of pores, they are often cylindrical and straight. The permeability of such pores is usually described by the Hagen-Poiseuille equation for flow through cylindrical tubes. If the membrane consists of a packed bed of particles, the specific permeability is derived from the Carman-Kozeny equation. Usually the pressure difference is applied externally to achieve filtration. During this process, solutes are separated from the solvent. These solutes accumulate near the membrane interface. The concentrations of these solutes are thus higher at the feed side than at the permeate side. This causes an osmotic pressure difference that reduces the effective pressure difference across the membrane. The simultaneous presence of both concentration and pressure driving forces across the membrane results in a mixed diffusive convective transport process.

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When both these forces are significant, both must be accounted for as well as their interactions. Irreversible thermodynamics provides a framework for the analysis of these complex interactions originally in the form of Kedem Katchalsky equations. More recently a modification of this approach was described by Kargol and Kargol9 who derived mechanistic equations based on the classical approach. These equations have the form: JV ˆ LP
P ÿ LP 
 and  P
P JS ˆ !D
 ÿ …1 ÿ †CL and represent both the volumetric (V) and solute (S) fluxes. These equations encompass four different processes; filtration, (LP
P) osmosis, (LP 
) solute  P
P) and where Lp;  and diffusion (!D
) and solute convection (…1 ÿ †CL ! are coefficients of filtration, reflection or rejection and permeation,  is the average concentration in membrane pores. respectively, and C If fluid alone is being transported across the membrane, then the first of the above equations will apply. The volume flux generated as a consequence of ultrafiltration can also be expressed in terms of concentrations by:   CW ÿ CP J ˆ K log n CB ÿ CP where CW , CP , and CB are the concentration of solute at the membrane interface, in the permeate stream and in the bulk respectively, and K is the mass transfer coefficient in the boundary layer. The degree of separation of a solute from the solution is expressed in a rejection coefficient, which is defined as: 1ÿ

CP CW

The rejection coefficient is dependent upon the molecular species, and its size relative to the pore of the membrane. Membranes are unable to discriminate between macromolecules that have sizes similar to one another, and for this reason membrane separation processes are unsuitable for the separation of soluble molecules whose molecular weight or effective sizes in solution (Stokes radii) differ by less than a factor of 6. When filtering low concentrations of macromolecular solutions by ultrafiltration, their osmotic pressures are low compared to the applied pressure and can be neglected. The osmotic pressure exerted by such a solution is generally in the form:
 ˆ AC ‡ A1 C 2 ‡ A2 C 3 where A, A1 and A2 are osmotic virial constants and C is the concentration of the bulk macromolecular solution. The coefficient A describes Van't Hoff's limiting

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law for the osmotic pressure which is applicable at very dilute concentrations, whilst A1 can be expressed in terms of the solute molecular weight. Ion transport is important in both biological and biotechnological applications. Ion transport across the membrane will be influenced by charge on the surface of the membrane (zeta or streaming potential) and the manner in which the ions are transported across the membrane. Transport across uncharged membranes may be characterised by consideration of the movement induced by the electrical field, diffusion and bulk convection. The movement resulting from each of these forces may not be in the same direction. The analysis of ion transport across charged membranes is more complex as the charge distribution on the membrane affects the concentration and potential distributions.

4.4

Blood material contact

Membranes and thin films used in therapeutic processes involve their repeated contact with biological fluids such as blood. For example, the treatment of renal failure by dialysis is generally undertaken three times weekly and each treatment session lasts 3±5 hours. Membranes have traditionally been considered as inert barriers, however, it is increasingly recognised that interactions occur between the material and blood components (Fig. 4.3). Thus, in therapeutic treatments

4.3 Biological pathway activation following contact with a material.

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this aspect of membrane performance becomes important. A primary requirement, therefore, of any such material is that the material should be biocompatible. The classical definition of biocompatibility is: `the ability of a material to perform with an appropriate response in a specific application'. This historic definition was based on the principle that a biomaterial has to perform and not simply exist, that it has to be associated with appropriate responses to ensure satisfactory performance. This definition recognises the fact that the response to a material will vary from one situation to another and that the appropriateness may vary. Furthermore, the definition allows a distinction to be made between biocompatibility and biological safety. The main difficulty with this definition is that the applications of materials in the clinical setting are varied, and there may be little commonality with the appropriateness of the responses. To account for this, the original definition has recently been revised such that the biocompatibility of a medical device that is repeatedly in contact with blood may be considered as `the ability of the device to carry out its intended function within flowing blood, with minimal interaction between device and blood that adversely affects device performance, and without inducing uncontrolled activation of cellular or plasma protein cascades'.

4.5

Biological events at the membrane and thin film blood interface

The magnitude of biological interactions is governed by the material's surface rather than its bulk characteristics, i.e., molecular interactions. Parameters of importance include hydrophobicity, which may be calculated using contact angle measurements, the presence of specific chemical groups (X-ray photoelectron spectroscopy and IR-spectroscopy), surface roughness (optical profiles and scanning probe microscopy) as well as the flow conditions at the time of the interaction.10 The underlying disease condition and drugs administered, may also play a role. A number of the responses that occur are in part a consequence of the blood recognising the material as non-self, and in part a consequence of the interaction between the material and cell surfaces. In addition cellular components of blood may also be involved, such as the activation of platelets with subsequent generation of thromboxane, the release of leucocyte proteinase enzymes and the generation of cytokines and reactive oxygen species (ROS).

4.5.1 Protein deposition or adsorption There is a common belief in biomaterials research that cellular interactions with natural and artificial surfaces are mediated through adsorbed proteins. The initial event that takes place on blood material contact is the deposition of proteins onto

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the surface.11 This deposition is rapid, occurring in a few milli-seconds and is governed by the interfacial properties of the proteins and the material surface. The importance of protein deposition stems from the fact that the deposited proteins exert a strong influence on subsequent cellular interactions with the surface. The process of protein adsorption comprises the following steps: 1. 2. 3. 4. 5.

transport towards the interface attachment at the interface eventual structural rearrangements in the adsorbed state detachment from the interface transport away from the interface.

The initial proteins deposited may be displaced by other proteins in the following sequence: albumin, immunoglobulins, fibrinogen, fibronectin, high molecular weight kininogen (HMWK) and Factor XII. This is termed as the Vroman effect. The adsorption of proteins may be considered to provide a degree of bioreactivity to the membrane or film surface. On the other hand, in the case of porous materials, such deposition inevitably modulates the diffusive characteristics of molecular species of interest. The adsorption of proteins mediates cellular adhesion to the adsorbed proteins. Primarily platelets are involved, and this sensitivity may be explained by the presence of receptors on the platelet surface (IIb/IIIa, Ib/IX) which facilitate such adhesion and are also involved in neutrophil activation. Many different techniques have been applied to study the molecular adsorption of proteins at interfaces. Several reviews have extensively discussed these techniques.12 The most commonly used techniques for protein adsorption kinetics studies include optical and spectroscopic techniques and electrical methods. Optical techniques include ellipsometry, variable angle reflectometry and surface plasmon resonance (SPR). Ellipsometry can be used with both transparent and non-transparent substrates whereas SPR must be used with a noble metal substrate.13 Spectroscopic techniques rely on the interaction of photons with the species present in the interfacial region to detect molecular events at the interface and include infra-red absorption (IR), Raman scattering, fluorescence emission circular dichroism (CD) and optical waveguide lightmode spectroscopy (OWLS). OWLS is a highly sensitive technique allowing real-time monitoring of protein-substrate interactions based on the measurement of the polarisability density (refractive index) in the vicinity of the waveguide surface. Since radioactive, fluorescent or other types of tagging are not required, measurements reflect the behaviour of unmodified proteins at the biomaterial interface. Other techniques include the use of radioactive isotope labelled proteins to quantitatively determine the extent of the protein adsorption or the spatial

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distribution on the surface of the material. Scanning probe microscopy techniques such as AFM provide dynamic imaging of real-time protein adsorption or static imaging of the adsorbed proteins. More interestingly, quantitative measurements of the strength of molecular interactions between the protein and the biomaterial are now possible.14 Protein-biomaterial adhesion force is measured with an AFM using a protein modified probe tip to analyse the actual interaction forces between blood proteins and the biomaterial surface.15 Predicting quantitatively the adsorption of a given protein at a well characterised surface under specified solution conditions remains a challenge.

4.5.2 Activation of the coagulation system Activation of the coagulation system occurs as soon as the blood leaves the vessel, and comes into contact with a non-physiological surface. Minimisation of coagulation may be by the use of drugs with antiocoagulant properties such as Coumarin, Warfarin, glycosaminoglycans, heparin and heparan sulfate. Plasminogen activators may also be used for the control of coagulation. When a foreign surface comes into contact with blood, the preotein deposition onto the surface includes Factor XII which becomes activated (Factor XIIa) and initiates the clotting or coagulation cascade. Fibrinogen is also present on the surface causing the attachment of platelets. The surface aggregated platelets encourage more fibrinogen to fibrin conversion by the production of additional thrombin and thus providing a positive feedback loop up regulating the coagulation system. The thrombotic potential of a material may be most readily assessed in a static experiment in which it is compared with a reference material, usually silicone coated soft glass. The thrombotic potential of the surface is generally evaluted by the timing of the development of a clot. These measurements may be combined with the determination of hemolysis.16 This type of approach, however, does not take into account flow across the surface which may influence the rate of clot formation or red cell damage. In-vivo testing of new materials poses problems, and often an indication of the behaviour of the material may be elicited by the consideration of ex-vivo tests. In such tests, a small volume of blood drawn directly from an animal of volunteer is passed over or placed into contact with the material and is then collected and analysed; considerable deviations from true in-vivo tests may be observed. When considering the behaviour of membranes used in extracorporeal procedures such as haemodialysis, additional issues are present. All therapeutic procedures utilise an anticoagulant, heparin, although less commonly low molecular weight heparin may be used. The adequacy of heparinisation during the procedure is generally measured by clotting time measurement. Of these, the measurement of activated clotting time is the

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simplest, however, it is only a broad non-specific indicator. A more accurate measure is the measurement of activated partial thromboplastin time (APTT). Heparin activity may be determined from the measurement of Factor Xa which is directly related to the amount of heparin present. Serial measurements provide an indicator of the loss of heparin due to adsorption or metabolism although in a clinical setting it is difficult to distinguish between these two. Commercially available kits are available for its determination. These measurements do not, however, provide any indication of the ability of the material to induce thrombus formation. A commonly used method for the measurement of this is by the measurement of thrombin antithrombin III complex.

4.5.3 Activation of the immune system The most thoroughly researched system activated following blood material contact has been the complement system. Activation of the human complement system following contact with a membrane or thin film is principally via the alternative pathway. It is initiated by the deposition of C3b on the material surface which with Factor B forms C3 and C5 convertases. These enzymes cleave the anaphylatoxins C3a and C5a from C3 and C5 by an autocatalytic process. Once in the circulation the C terminal arginine is removed and C3a des Arg and C5a des Arg are formed. These fractions are generally measured when membrane induced complement activation is studied. The cleavage of C5 by C5 convertase results not only in the production of C5a but also of C5b which initiates a macromolecular complex of proteins, the membrane attack complex, formed from C5b, C6, C7, C8, C9. (C5b-9). Complement products are an important mediator of the inflammatory response produced by membranes and are important markers for the biocompatibility of such materials. A wide range of clinical sequalae are associated with the generation of C3a, C5a and C5b-9 and virtually every blood cell type responds either directly via receptors, or indirectly via secondary mediators to these complement fractions. Implantation of devices in contact with blood may be associated with interactions between the implanted material and the host immune system. The aberrant state of monocyte and T-cell activation resulting from these host/device interactions is accompanied by two parallel processes, first, selective loss of Th1 cytokine producing CD4 T-cells through activation-induced cell death and secondly, unopposed activation of Th2 cytokine producing CD4 T-cells resulting in B-cell hyper-reactivity and dysregulated immunoglobulin synthesis through Th2 cytokines and heightened CD40 ligand-CD40 interactions. The net results of these events is that the patient develops progressive defects in cellular immunity and is at risk of infection, or allosensitisation.

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In characterising complement activation by biomaterials, older publications dealing with complement activation refer to a less sensitive measure, total complement (CH50), to quantify activation. As discussed above, the primary event is the interaction of the C3 molecule with the material surface. This is followed by a complex series of cascade reactions during which C3a and C5a are released. Upon release into the circulation there is removal of the C-terminus arginine from C3a (C3a des Arg). Such modification abolishes anaphylotoxin activity but does not affect the platelet-stimulating activity of the peptide or release of serotonin.17 The circulating C3a des Arg levels may be measured by commercially produced ELISA kits, but currently no measurement of C5a des Arg is possible. Commercially produced ELISA assays are also available for the quantification of the membrane attack complex (C5b-9).

4.5.4 Cellular activation Leucocytes White blood cells are made up of neutrophils, eosinophils, basophils monocytes and lymphocytes. The principal functions of neutrophils and monocytes are their ability to recognise foreign material and in the body's response to infection. They are attracted to sites of inflammation and phagocytose foreign particles. Following phagocytosis, neutrophils degranulate. Macrophages share the ability of neutrophils to phagocytose foreign materials, but can ingest larger particles and deal with a greater number of particles. In therapeutic procedures such as haemodialysis, plasma separation and membrane oxygenation, following initial contact of the blood and foreign surface, a profound transient neutropaenia occurs. The neutrophils are known to aggregate in the lung vasculature, and it has been shown that the magnitude of neutropaenia is related to the degree of complement activation, however, the molecular mechanisms governing this feature of blood material contact remain incompletely elucidated. It is likely that the activation of neutrophils produced by C5a induces changes in the expression of leucocyte endothelial adhesion molecules as neutrophils have receptors for C3b, C5a, and integrins such as MAC 1 (CD11b/CD18) responsible for adhesion, the modulation of which leads to cell adhesion to the endothelium, and pulmonary sequestration. Measurement of such molecules may be by the use of ELISA kits or by the use of fluorescent antibody cell sorting (FACS) analysis. Neutrophil activation is also associated with superoxide production which can be measured following material contact. Granulocytes (polymorphonuclear neutrophils) release proteinases such as elastase, lactoferrin and myeloperoxidase which may be measured by ELISA kits.

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Platelets Circulating platelets adhere to proteins deposited on the material surface. Changes in circulating platelet numbers may be elucidated by serial measurements. The deposition, however, is not only dependent upon the material but also on the flow conditions. The deposited platelets on the material surface release their granule content. Granules may be alpha or dense. Alpha granules contain platelet factor 4 (PF4) and B thromboglobulin (BTG) Both may be used as markers of activation, however, in a clinical setting other factors need to be considered, namely the clearance of PF4 from the circulation by endothelial binding. Heparin may also induce the release of PF4 from endothelial cells. Release is governed by the material surface as well as circulating factors such as thrombin and thromboxane A2. The dense granules contain ADP, ATP pyrophosphate and serotonin. Kits are available for the measurement of PF4 and BTG. The up-regulation of surface markers on platelets such as CD62 following material contact can be quantified by FACS. Platelet neutrophil interactions Platelet leucocyte co-aggregate formation has been implicated in the pathogenesis of thrombosis and inflammation, and observed during haemodialysis.18 Such interactions may be of relevance in both haemostasis and inflammatory processes. In haemodialysis, the increased formation of platelet leucocyte micro-co-aggregates is thought to be related to a primary platelet activating mechanism that involves P selectin (CD62P) a marker of activated platelets as well as CD15s (the sialyl-Lewis x molecule) a selectin ligand and it is possible that the CD62P/CD15s interaction seen during haemodialysis represents the first stage of leucocyte margination. Monocytes Monocytes are activated during blood membrane contact. C5a induces mRNA IL- and TNF transcription and primes the cells for the translation of cytokines following further stimulation. This further stimulation may be endotoxin entering the bloodstream from the dialysis fluid across the membrane, or by the direct contact of the monocytes with the membrane itself. Eosinophils Eosinophils have anti-histamine properties and congregate around sites of inflammation. Normal blood contains between 1±6% eosinophils, and elevated numbers are known to be present in people with allergic conditions such as hay fever. Eosinophilia is known to occur in haemodialysis patients; its presence has

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been associated with allergy to dialysers and exaggerated activation of complement during HD. Its etiology, however, remains unknown. Erythrocytes Red cells are remarkably robust, however, mechanical damage may occur during extracorporeal circulatory processes arising from the blood pump or through rupture of the cell membrane by coming into contact with granular material. Lysis of red cells results in the release of ADP which is a potent mediator of platelet aggregation.

4.6

Improvement of biocompatibility

Improvement of the biocompatibility profile of a membrane material is generally achieved by the alteration of its bulk characteristics by alteration of the manufacturing process or the use of polymer blends. Increased understanding of the molecular mechanisms involved is resulting in better understanding of the role of material related properties influencing the molecular mechanisms and forms the basis of the development of materials with minimal bioreactivity. Controlled chemical modification can lead to the development of predictable biocompatibility profiles. For example, hydrophilic domains on the material surface in polymer blends have a stimulatory part on the complement activation potential of the material, but play little part in determining the material's ability to activate platelets. On the other hand, hydrophobic domains show a reduced influence on the activation of the complement system, but stimulate platelet adhesion.19 Plasma modification of the membrane surface may be a useful way to improve the biocompatibility of polymer membranes by utilising high-energy electrons, ions, atoms, radicals and excited molecules produced by electric discharge. The plasma generates chemically reactive species from otherwise inert molecules (at relatively low temperatures) that may be deposited on the membrane surface (changing its chemical composition) thereby affecting its biocompatibility (Table 4.3).

4.7

Conclusion

Although membrane separation processes are widely used in many major industries a number of unresolved technical requirements remain. These include the availibity of pre-evaporation membranes for organic-organic separations, oxidation-resistant membranes for reverse osmosis and ultrafiltration foulingresistant membranes. Electrodialysis is an established membrane separation process for the desalination of brackish waters which has changed little in the last ten years. Membranes with better temperature stability and spacers with

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Table 4.3 Biomedical applications of plasma modified polymer surfaces (adapted from Scott, K. 1995, Handbook of industrial membranes, 1st edition. Elsevier Science Publishers Ltd) Gases (or monomers)

Polymers

NH3 (or N2 and H2)

Polypropylene (PP) Poly (vinylchloride) (PVC) Polytetrafluoroethylene (PTFE) Polycarbonate (PC) Polyurethane (PU) Polymethylmethacrylate (PMMA)

Heparin bonding for improved blood compatibility

Hexamethyldisiloxane (HMDS) C2H4 + N2 Alllene + N2 + H2O

Poly (ethylene terephthalate) (PET) Silastic (SR) Polysulfone (PS)

Improved blood compatibility

C2H4, allene, styrene, acrylonitrile, C2F4, C2H3F, C2F3Cl, C2H3Cl

Polystyrene (PSt) SR

Improved tissue compatibility

C2H4, C2F3Cl styrene

SR

Improved tissue compatibility

C2H2 + N2 + H2O

PMMA

Modified contact lens wettability by proteins

improved flow distributions can produce incremental improvements in brackish water desalination systems. A number of novel technologies are also being developed, including magnetic separations, electrically driven systems and liquid membrane systems. In membranes used for clinical applications an ongoing focus is improvement of biocompatibility. New manufacturing techniques involving a more defined pore distribution have also been introduced.20 Tissue engineering offers a number of exciting opportunities for the treatment of a variety of medical conditions. The membranes and thin films provide an ideal artifical support structure for the development of bio-hybrid devices. Such devices are available for both kidney and liver support.21,22 Micro-encapsulation offers the potential to encapsulate the islets of Langherhans for the treatment of diabetes. Such encapsulation may also be used for affinity dialysis for the removal of HIV and toxic viral proteins from blood.23

4.8

References

1. Hoenich NA, Woffindin C, Brennan A, Cox PJ, Matthews JN, Goldfinch M (1996) A comparison of three brands of polysulfone membranes. J Am Soc Nephrol.; 7(6): 871±6.

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2. Vanholder R, Smet RD, Glorieux G, Dhondt A (2003) Survival of hemodialysis patients and uremic toxin removal. Artif Organs.; 27(3): 218±23. 3. Colton CK (1987) Analysis of membrane processes for blood purification. Blood Purif.; 5(4): 202±51. 4. Sauer IM, Goetz M, Steffen I, Walter G, Kehr DC, Schwartlander R, Hwang YJ, Pascher A, Gerlach JC, Neuhaus P (2004) In vitro comparison of the molecular adsorbent recirculation system (MARS) and single-pass albumin dialysis (SPAD). Hepatology; 39(5): 1408±14. 5. Klingel R, Fassbender T, Fassbender C, Gohlen B (2003) From membrane differential filtration to lipid filtration: technological progress in low-density lipoprotein apheresis. Therap Apher Dial.; 7(3): 350±8. 6. Wickramasinghe SR, Goerke AR, Garcia JD, Han B (2003) Designing blood oxygenators. Ann N Y Acad Sci.; 984: 502±14. 7. Isayeva IS, Kasibhatla BT, Rosenthal KS, Kennedy JP (2003) Characterization and performance of membranes designed for macroencapsulation/implantation of pancreatic islet cells. Biomaterials.; 24(20): 3483±91. 8. Humes HD, Weitzel WF, Fissell WH (2004) Renal cell therapy in the treatment of patients with acute and chronic renal failure. Blood Purif.; 22(1): 60±72. 9. Kargol M, Kargol A (2003) Mechanistic equations for membrane substance transport and their identity with Kedem-Katchalsky equations. Biophys Chem.; 103(2): 117±27. 10. Spijker HT, Graaff R, Boonstra PW, Busscher HJ, van Oeveren W (2003) On the influence of flow conditions and wettability on blood material interactions. Biomaterials.; 24(26): 4717±27. 11. Ramsden JJ (1995) Puzzles and paradoxes in protein adsorption. Chemical Society Reviews; 24: 73±78. 12. Ramsden JJ (1994) Experimental methods for investigating protein adsorption kinetics at surfaces. Q Rev Biophys.; 27(1): 41±105. 13. Voros J, Ramsden JJ, Csucs G, Szendro I, De Paul SM, Textor M, Spencer ND (2002) Optical grating coupler biosensors, Biomaterials; 23: 3699±3710. 14. Kidoaki S, Matsuda T (2002) Mechanistic aspects of protein/material interactions probed by atomic force microscopy. Colloids and surfaces B.; 23: 153±163. 15. Kidoaki S, Matsuda T (1999) Adhesion forces of the blood plasma proteins on selfassembled monolayer surfaces of alkanethiolates with different functional groups measured by an atomic force microscope. Langmuir; 15: 7639±7646. 16. American Society for Testing Materials (1997) ASTM F 756-93. 17. Polley MJ, Nachman RL (1983) Human platelet activation by C3a and C3a des-arg. J Exp Med.; 158(2): 603±15. 18. Gawaz MP, Mujais SK, Schmidt B, Blumenstein M, Gurland HJ (1999) Plateletleukocyte aggregates during hemodialysis: effect of membrane type. Artif Organs.; 23(1): 29±36. 19. Deppisch R, Gohl H, Smeby L (1998) Microdomain structure of polymeric surfaces ± potential for improving blood treatment procedures. Nephrol Dial Transplant.; 13(6): 1354±9. 20. Ronco C, Bowry S (2001) Nanoscale modulation of the pore dimensions, size distribution and structure of a new polysulfone-based high-flux dialysis membrane. Int J Artif Organs.; 24(10): 726±35. 21. Humes HD, Weitzel WF, Fissell WH (2004) Renal cell therapy in the treatment of

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patients with acute and chronic renal failure. Blood Purif.; 22(1): 60±72. 22. Patzer II JF, Lopez RC, Zhu Y, Wang ZF, Mazariegos GV, Fung JJ (2002) Bioartificial liver assist devices in support of patients with liver failure. Hepatobiliary Pancreat Dis Int.; 1(1): 18±25. 23. Tullis RH, Duffin RP, Zech M, Ambrus JL (2003) Affinity hemodialysis for antiviral therapy. II. Removal of HIV-1 viral proteins from cell culture supernatants and whole blood. Blood Purif.; 21(1): 58±63.

5 Stable use of biosensors at the sample interface J F G A R G I U L I , University of London, UK, A G I L L and G L I L L I E , University of Manchester, UK, M S C H O E N L E B E R , University of London, UK, J P E A R S O N , University of Manchester, UK, G K Y R I A K O U and P V A D G A M A , University of London, UK

5.1

Introduction

Biosensors provide a high degree of elegance in regard to their simple juxtaposition of a bioreagent and a transducer function. To work properly, there must be a direct alignment of a functionally responsive biolayer and a transducer element, which is able to directly extract the binding information resulting from the encounter with the analyte. There is a difficulty associated with such a simple, structurally inflexible combination, due to the fact that optimisation is limited as compared with say the use of a liquid phase bioreagent with its attendant optimised solution parameters of pH, pI, concentration and reagent additives. However, the net result is a solid-state monolithic structure with the potential to perform analyses and to operate potentially in optically opaque samples. In biomedicine, the latter capability holds considerable advantages. Indeed, most if not all biofluids contain colloidal materials that are liable to render the sample optically opaque or, at the very least, to induce a certain amount of light scattering. Furthermore, the majority of clinical sample assays rely on absorbance techniques. Therefore, it is necessary to have access to biosensors that would perform reliably even in opaque samples. Nonetheless, there are still a few drawbacks associated with the use of biosensors in vivo. Although the established biosensor systems operate on the `macro' scale and have seen varying degrees of clinical exploitation,1 a key reason for their limited introduction into the application domain has been the rapid alteration of the biosensor interface through the surface activity of the colloidal elements of any unmodified biological sample. This problem is also intricately linked to that of biocompatibility of the exposed surface of the biosensor in direct contact with either living tissues or the physiological fluids present in the body.

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5.2

Biosensor limitations

5.2.1 Biosensors and bioresponse compromise Whilst the bulk amount of protein, colloid and cell transfer to a `clear' biosensor surface may be relatively low in amount, its diffusional barrier effect on the continued flux of analyte to a responsive biolayer surface will immediately be registered as a reduced biosensor response. The function of any flux dependent biosensor, unless it is merely a qualitative registering device, will be affected sufficiently that accuracy and precision of measurement will be lost, and neither may be recovered simply by recalibration. The devices therefore likely to be most affected will be those using degrading enzymes rather than where a true binding equilibrium is approached, e.g., antibody, lectins, receptors and DNA/RNA. In the case of the latter group, the only effect should be on the rate of approach to equilibrium not on its final value, unless of course the fouling biolayer affects solute partitioning. This is possible in principle if, say, the fouling layer is a charged colloid, and the analyte target is also charged; electrostatic forces are then liable to come into play to either partition in or partition out a given analyte. Exclusion, furthermore, is possible if the analyte target is a macromolecule, in which case it may have limited access to the affinity surface of the biosensor through the limited permeability of the colloid fouling layer. The underlying transducer of the biosensor, whilst not directly affected by surface colloidal deposits, may register the presence of these non-specific elements through its detection domain. Thus, the evanescent wave of an optical wave-guide or SPR system will respond equally to non-specific binding as to specific affinity interactions. Passivation of the transducer element is possible with some surface-active crystalloids. Thus free amino acids and thiol-containing molecules that are surface active can distort and depress the catalytic behaviour of a Pt working electrode used for redox dependent biosensors.

5.2.2 Biosensors and the selectivity compromise Whilst the biological component of a biosensor has accepted selective properties, and is de facto the driver for biosensor development, the underlying transducer, whether based on electrochemical, potentiometric, optical or microgravimetric principles, is vulnerable to a false positive response due to the surface activity of analogue species of either the target molecule or a molecule that is part of the transduction cascade. The most potent expression of the problem is where an electrically polarised noble metal electrode is used to detect the H2O2 product of an oxidase enzyme catalysed reaction in an enzyme electrode. At the typical polarising voltage of +0.65 V vs. Ag/AgCl, numerous species in biological solution are simultaneously oxidased, and false positive responses therefore result.2

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5.2.3 Interfacial problems at microfabricated biosensors The issues of direct biosample interfacing of biosensors applies to all biosensors irrespective of length scale. So whilst the x-y plane architecture may be geometrically precise, formulated in a much more reproductive manner using MEMs and other microforming technologies, the biological response of the host sample reacts in rather similar ways, with adverse buildup along the z axis, i.e., normal to the sensor surface. One rider to this equivalence of macro/microsensor outcomes is where response is flux (continuous diffusion) dependent; a sufficiently small sensing microsurface will have spherical and not a planar diffusion based supply of analyte and is thereby less affected by external variables such as fouling. There is also evidence that a microstructure may set up a lower intensity tissue reaction thus leading to reduced surface fouling.

5.3

Biocompatibility

The term `biocompatibility' covers the whole range of interactions that exist between a biomaterial implant and its biological surroundings, as well as the orchestrated sequence of responses the body invokes to essentially reject that implant as non-self. In the case of invasive in vivo monitoring, an electrochemical biosensor requires intimate, direct contact with the sample matrix in order to function properly, notwithstanding the intensity of the body's reactive response to its constituent materials. As a consequence, the observed performance of the biosensor is highly vulnerable to the local accumulation of surface-active agents from the body such as cells, proteins and other less well identified constituents such as colloidal and lipid aggregates. This accumulation of biological compounds on the surface of the sensing device is known as `biofouling', and will with time alter and degrade the biosensor response and performance. Furthermore, biosensors are not bio-inert, not only because of the active redox components they may incorporate but also because of the polymeric materials, as well as the coated or uncoated metal and carbon electrode interfaces they present to the living tissues. In these tissues, they provoke a high intensity local inflammatory reaction simply because they inhabit a wound site and, in blood, they are a nidus for surface coagulation with the attendant threat of local microand later macro-thrombi. The danger of thrombus dissemination in the vasculature, thromboembolism, constitutes a particular concern because of the possibility of considerably wider distribution of tissue damage. There are many examples of implants leading to thrombosis, embolism and death. While the risk might be acceptable for a life-saving device such as a vascular stent, such risk cannot be tolerated when a diagnostic device such as a biosensor is being used. Even if an implanted material is reputed to have no interaction with a biomatrix and is therefore believed to be bio-inert, some interactions at a microscopic scale still occur. A completely biocompatible material does not

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exist. As a consequence, all available materials provoke, to some extent, a biological response, adverse or not, that will impact on overall sensor performance. Indeed, even air entrapped within a tissue will ultimately provoke a local body reaction. Additional bio-effects, beyond local surface phenomena, that need to be taken into account when designing a biosensor for clinical use are carcinogenicity, mechanical stability, immunogenicity, chemical stability and biomechanical compatibility with the local soft tissue. Also, sensors may need to operate in vivo for quite different periods of time, and this must be taken into account when assessing the tolerance limit for biocompatibility. Admittedly, in contrast to the carefully assessed quantitative analytical behaviour of sensors in vitro, little progress has been made in regard to in vivo performance standards and acceptability. Also, the actual limits of tolerance for the degradation of function, resulting from the body's response, are still poorly understood and known. In vivo biocompatibility can be regarded as a hybrid between biomaterials and biosensor research, which deals with both the specific, as well as the complex interactive and cumulative effects of sample matrix constituents upon sensor function and operational life-time.3±5 For all sensors, understanding and predicting the in vivo biological response still poses a great, unmet, challenge. The crucial point is that, in contrast to conventional biomaterials, the surface deposition of the body's diffusive and cellular biocomponents leads to degradation of function, observed within minutes and hours rather than days and months. Whilst the device continues to function, its value as an accurate and precise quantitative system is lost. A hierarchy of biological interfacial phenomena exist, which, though subclassified relatively easily, are difficult to unravel in relation to their complex dynamics and as concerted interactive phenomena. For a long time researchers attempted to avoid the bioresponse entirely via the `mythical' concept of the totally bio-inert implant, leading later to the emergence of the idea of functional biocompatibility.6 In this latter case, a primary issue is not so much the total avoidance of the bioresponse, but rather achieving sufficient control over its adverse effects on the biosensor performance. The definition of biocompatibility is itself an approximation of multiple concepts. However, it is based on the expectation that certain types of materials will be able to provoke just a limited body response. No matter what the material deposited and accumulated at a sensor surface may be, it is whether or not a degradation of response occurs that really matters. The counterpart to this is whether an implant material poses a threat to the patient or not. Both immunological and toxicological dangers exist, for example, through the leaching of additives within covering polymers, the release of polymer degradation products, of metal/metal oxide particulates or of carbon electrode constituents. All are capable of triggering an early inflammatory and toxicological response, but the dangers of long-term effects, including

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carcinogenicity are largely uncharted. Again such dangers may be acceptable in the instance of acute life-saving or life quality enhancing implants such as heart valves, aortic grafts and pacemakers, but in the case of a measurement modality such as glucose monitoring, any such dangers have to be negligible. A biosensor can be regarded as a composite material, usually made up of metallic, polymeric and biological components. While it constitutes a minor burden as an implant, as it exhibits trivial mass and volume in relation to the body, it is however more likely to present a multiple combination of potential leachables, which, together with the reaction products of the biosensing reaction, could have significant local effects. Underlining all this is the fact that any release of antigenic protein poses special dangers as an immunogenic trigger, especially if a biosensor is to be implanted repeatedly. The biocomponent (e.g. enzyme) may itself be associated with an undefined constituent due to the presence of low level impurities, as in any bioagent, rendering traceability difficult. A practical expression of this comes from the routine use of bovine serum albumin (BSA) as a crosslinking matrix for glucose oxidase in the early days of research.7 Due to actual and theoretical dangers of the bovine spongiform encephalitis (BSE) agent, such a protein source for enzyme immobilisation is no longer acceptable. In view of the heterogeneity of both sensor internal structure and of the multiple types of surfaces presented to the biomatrix, the body's response maybe quite different over the active sensing regions of the device vs. the support regions. This makes it difficult to determine a precise structure ± biocompatibility linkage.8 Materials considerations are also relevant to the duration of operation envisaged. Short-term monitoring in, say, the critically ill patient, demands quite a different level of materials requirement to the more robust, mechanically resistant devices that, moreover, would have to be well tolerated by the patient over the long term, especially during ambulatory monitoring. Polymer encapsulant degradation and metal corrosion set a limit to the long-term implantation of the sensor in the patient. Furthermore, not only are the body's responses cumulative but a rigid device, mechanically incompatible with local soft tissue, will not be well tolerated, due to the stimulation of local pain sensors. Over time these released constituents, due to the immune response, will be upregulated; so will be the effects of unavoidable products of enzyme reactions and redox centre/mediator components. Wherever a foreign body is lodged in the tissue or the blood compartment, it can also serve as a focus for infection and poses a problem over the long term. Microbial films may form on the surface and these have a powerful way of resisting antibiotic assault. Such films are known as significant contributors to hospital infection rates.9 The situation is compounded if a device is only partially inserted as with, say, a percutaneous glucose needle sensor. These are typically implanted in vascular, partly fatty tissue to only a 10±15 mm depth. Whilst less invasive, they provide a contact pathway for externally derived

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microorganisms. Important in this regard is the fact that the skin surface has its own microbiological flora and cannot be fully sterilised. Diffusible, low molecular weight solutes in biological fluids provide a nominally stable solution environment in vivo, given that this internal environment is designed for physiological stability. Certainly, near-neutral pH conditions normally prevail, and the concentration variation of background electrolytes is relatively small, at least with regard to the major ions such as sodium, potassium and calcium. As an indicator of the total solutes, osmolality in blood ranges quite narrowly between 280±295 mOsm/kg. With the inherent stability of the enzyme within an electrochemical biosensor and its reduced dependence upon pH and background ionic changes due to the immobilisation, one would expect minimal background solute effects upon bioelectrochemical sensor function. However, even small variations will cause some analytical imprecision, especially as sample dilution or other specimen manipulation are precluded in contrast to bench top analysis. The above considerations have been a basis for the use of, say, sampling tissue ultra-filtrate for glucose monitoring via mechanical suction through permeabolised skin.10 However, low molecular weight solutes may still influence working electrode response through adsorption, passivation and then direct surface activity. Such effects may be difficult to identify when conducting measurements in a complex medium such as blood or tissue combining macroand microsolute influences. The former are only external surface active; the latter are able to permeate the entire device. The adsorbed solute can induce degradation of metallic components, as described for electrodes used in electrical stimulation, where increased ionic release and accelerated corrosion have been reported11 with a further facilitating influence due to surface active proteins.12 The electrochemical reaction can also contribute to electrode dissolution, which is of relevance to long-term implantation.13 Beyond standard cylindrical geometries and wire-type devices, emphasis is shifting to MEMs-based devices. However, some implantable sensors based on MEMs constructs are vulnerable to hydration and water ingress. As a consequence, rigorous packaging and encapsulation are needed for long-term function as well as over the short term. Typically, in order to avoid leakage currents and extraneous responses, inorganic or polymeric packaging materials are required.14,15

5.3.1 Protein constituents In terms of the total colloid biomass, proteins comprise the most important surface-active constituents of biofluids other than viable (and dead) cells. Total concentration in plasma is around 80 g/l. Proteins have a special ability to adsorb physically at solid/liquid interfaces and through consequent denaturation to form relatively adherent, immobilised phases, showing only partial remodelling or

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recycling between the bulk sample and surface. The adsorbed proteins thus undergo both conformational and orientational changes at the surface. As well as progressively increasing in depth with time, through deposition of solution originated proteins, the entire protein `gel' layer grows unpredictably, modulated by both environmental perturbation (convective, shear, pressure) and inherent structural remodelling, especially evident near the growth surface.16 Again, the key issue is that in contrast to conventional biomaterials, surface deposition effects on a new electrode surface are rapid and accentuated. Whilst the device continues to function, its value as an accurate and precise quantitative system is almost immediately lost, a particular concern in the case of glucose monitoring given the strict control demanded in diabetes management. Protein deposition occurs within seconds of contact,17 but inherent film composition and individual component abundance are conditioned by material surface energy (hydropholicity/hydrophilicity), charge, charge density, surface profile and roughness, degree of molecular ordering, pendent group flexibility and crystallinity. In the case of a polymeric interface, the presence of trace components, plasticiser content and minor degrees of surface oxidation all drive particular types of protein adsorption. Albumin, in particular, is a dominant constituent at the interface, and outcompetes other proteins in plasma such as immunoglobulins, fibrinogen and kininogen. Its presence at a surface, furthermore, helps to reduce the deposition of other surface fouling components including cells, and it is regarded as a passivating protein. The dynamics of the adsorption, however, appears to be different at different surfaces. For example, on a hydrophobic surface, adsorption is a single-step process, whereas on a hydrophilic surface, it involves two steps. The resultant in either case, though, may be the formulation of an equivalent layer with similar passivating behaviour.18 This new view of protein deposition reduces the importance of the primary surface in driving biocompatibility outcomes. Ongoing competition between proteins for the surface leads to the remodelling of the deposited surface layer and, in the context of, say, fibrinogen adsorption from plasma, its transient surface prevalence and subsequent partial displacement (the Vroman effect) represents the outcome of the finite number of available sites on the surface.19 In previous studies, surfaces have been alkyl modified in order to promote albumin deposition, and therefore to reduce platelet adhesion.20 More recently, immobilisation of highly hydrophilic polymers such as poly(ethylene oxide) (PEO) has proved especially valuable in resisting protein deposition.21 The equivalent has also been achieved, for example, through glow discharge deposition using tetraethylene glycol dimethyl ether.22 More classically, heparin immobilised at surfaces has proved effective in resisting a special type of surface deposition, the binding and activation of coagulation components. In this respect, surface bound heparin has been used to good effect at an intravascular oxygen catheter sensor.23

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Heparin works through its anionic sulphonate and aminosulphate groups,24 so attempts have been made to sulphonate artificial polymers, e.g., polystyrene and polyurethane.25 Also, the combined effect of PEO with end-attached sulphates to give heparin-like properties, and the added benefits of PEO flexibility has led to an improved thrombo-resistant polyurethane.26 Polyurethane is already frequently used as an outer membrane barrier for glucose sensors, so this may have a direct application when used in a bioelectrochemical sensor. The fundamental drawback of an enzyme-based biosensor is that it requires continuous substrate flux to the enzyme layer for an ongoing response. This flux has to be purely substrate concentration (gradient) dependent, unperturbed by any newly formed diffusion barrier in, on or around the biosensor itself. The first of the membrane-packaged devices developed to control such adventitious biolayer diffusion limitation was the Clark pO2 polarographic electrode where an external gas permeable membrane served to protect the working electrode from protein and colloid deposition.27 For surface anticoagulation, a primary need is for the prevention of the binding of factor XII, as this is a key trigger in the initiation of the coagulation cascade and eventual deposition of the crosslinked fibrin mat at the surface. An intrinsic pathway and an extrinsic coagulation pathway, the latter being induced by tissue-derived factors, may combine to create an accelerated cascade of fibrin deposition (Fig. 5.1). However, a parallel system of surface-active proteins has been rather neglected. This is the complement system, which undergoes an ordered, sequential response to an artificial surface, eventually to trigger the production and the release of flammatory mediators from white cells.27 This complement activation takes place through one of two pathways, the classical and the alternative. The classical complement pathway is initiated by antigen-antibody complexes, by crystals or bacterial and virus surfaces if antibodies are absent, or by complexes between positively and negatively charged molecules such as those between heparin and protamine. The alternative pathway is not triggered by immune/antibody complexes, but can be initiated by any foreign material introduced into the body, including a biomaterial, lipopolysaccaride, polysaccharide, or bioorganism. It is activated by surfaces with particular chemical characteristics allowing fragment C3b of a larger protein C3 at the surface to initiate the assembly of an amplification system, C3 convertase, at the surface28 for further C3b deposition (Fig. 5.2). With regard to specific sequences, C3 convertase cleaves C3 to generate C3a, and a further fragment C3b. In the nascent state, the latter binds to the surface and augments the convertase enzyme further to amplify C3b deposition on the surface. A further reaction leads to the cleavage of available C5, with the production of a C5a fragment and also recruitment of a C5±C9 sequence of effectors and associated inflammatory changes. It is clear that substantial complement activation can lead to major organ dysfunction, though admittedly

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5.1 Surface coagulation cascade initiated either by surface contact (the intrinsic pathway) or by tissue factor (TF, the extrinsic pathway). The two pathways are eventually converting, forming a fibrin clot due to activation of thrombin on fibrinogen. Factor XIII will eventually convert fibrin clot into insoluble fibrin gel.

only following large scale blood/surface interactions as in haemodialysers.29 However, this complex cascade also requires to be considered as a possible contributor to local events at the biosensor surface. Surface amines and hydroxyls, in particular, react with C3 to form complexes, though there is uncertainty whether these are covalent bonds or whether electrostatic and hydrophobic interactions are important.30 Outcomes with regard to systemic effects also need to be unravelled.31 Local effects of relevance to a bioelectrochemical sensor might be abated through control of the surface presented, perhaps through surface heparinisation.32

5.3.2 Blood interfacing Blood consists of about 55% plasma by volume with about 45% solid particles including proteins and cellular elements, especially red blood cells, but with a

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5.2 Surface induced complement activation leading to accelerated complement protein deposition at a surface expressing amino/hydroxyl groups. Side cascade generation of C5±C9 is also indicated.

small white blood cell and platelet volume contribution. Red blood cells contain mainly haemoglobin, carry oxygen to the peripheral tissues and normally do not leave the circulation. White blood cells, however, are able to leave the vascular system to move towards any active disease tissue focus from, say, that due to microorganisms intrusion through a foreign body. The induced tissue disruption is `sensed' very acutely, no matter how localised. Additionally, platelets operate as a natural blood containment system, typically by forming a mechanical microplug at a vascular injury site, eventually producing a defined clot to contain blood in the circulation. The cellular elements of blood play a complex, cooperative role in the maintenance of the surface coagulation process initiated by soluble coagulation factors (vide supra). However, the deposited protein layer from blood at the sensing surface changes the material interface considerably, and it is this layer over which the cellular elements then begin to accumulate. Initially, the most important of these are platelets, which normally circulate in the blood in an inactive state, but are also the most labile of all the formed elements and therefore the most difficult to evaluate in physiological studies. Once they come into contact with a foreign surface, they show strong adhesion. This spreading and aggregation behaviour, associated with the release of intra-platelet adenosine diphosphate (ADP), promotes secondary platelet aggregation and the creation of (adherent) thrombus. In the final analysis, the level of initial coagulation protein deposits at a surface are seen to correlate with platelet

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surface activity and therefore, both condition the blood compatibility of a given material.33 Quite apart from any surface effects, platelets are exquisitely sensitive to shear force. Both the magnitude of shear and the overall blood flow profile near a surface, through influence on platelet transport, can completely override surface affinity interactions. Moreover, turbulent conditions and their associated high shear are particularly able to trigger platelet activity and augment surface delivery.34 The platelet delivery process is further accentuated by red cell interactions and local fluid entrainment around red cells, promoting surface collisions.35 The net negative charge on a platelet, due to surface sialic acid groups, is important to surface attachment. However, surface receptor interactions are also a powerful determinant, such as the receptor mediated attachment of platelets to fibrinogen. Following adhesion, platelets degranulate with the release of ADP. Then, a host of other bioactive components are able to promote further platelet activation and the eventual creation of an adherent micro-thrombus. Incorporation into the thrombus of the other cellular components of blood, including white cells and red cells (both influenced by blood flow and pressure gradients), leads to thrombus growth. As a consequence, local blood flow is distorted near the originally smooth surface, further stimulating the growth of thrombus and aggregation of blood components of all types. The effect on a sensing surface is the reduction of solute transport to that surface and the local consumption of oxygen/glucose, depending upon the metabolic state of the cellular aggregates, thereby inducing changes on the locally measured parameters. It becomes especially difficult to characterise material-induced artefactual influences upon a sensor, let alone to model them. Therefore, their subtraction from true responses in vivo becomes problematic, demanding at the very least a frequent in-vivo calibration regimen. It may be that, in the future, the quantity and nature of a surface coagulation phase can be measured using an independent technique. Impedance measurements promise in this regard to allow at least for baseline drift36 but, in reality, the end result at present is the highly unsatisfactory need to recalibrate frequently. The imposed structure of an intravascular electrode disrupts blood laminar flow patterns and can be an indirect cause of surface thrombus formation. Experience with thrombolic events and clinical use of intravascular catheters indicates a relationship between vessel cross-section diameter and catheter size.37 Complicating the assessment of protein deposition, coagulation/ complement activation and platelet retention,38,39 surface irregularity and surface microdepressions over a device may promote thrombus formation.40,41 A smooth surface is an advantage, but, probably, surface anticoagulation is a more effective basis for reducing thrombus formation. There is a clear lesson here for the combination design of smooth, haemodynamically acceptable glucose sensor surfaces and their surface modification.

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Heparin is a natural anticoagulant in blood, and is the most frequently used bioactive surface ingredient used to reduce clotting. The protective action of heparin is based upon its stimulatory effect on antithrombin via an heparinantithrombin complex, though this may be countered by the fibrin interactions of thrombin.42 Surface heparinisation of membranes has allowed more reliable oxygen monitoring.43 Catheter heparinisation has also proved to be of general effectiveness.44 Covalent binding of heparin, though permanent, leads to a more rigid attachment which can reduce effectiveness so bridging groups are an advantage. Also, depending upon the required duration of sensor operation, heparin leaching from a porous membrane or other reservoir could lead to a more potent anticoagulation surface. Overall, heparin can play a major role in reducing the effects of coagulation, but it is not the complete solution to the problem hoped by some. Alternatives to heparin include low molecular weight anionic analogues and, in particular, poly(ethylene oxide) (PEO).45 Furthermore, the use of copolymer structures rather than surface attachment may provide a family of new blood-stable materials in the future.46

5.3.3 Tissue interfacing Subcutaneous tissue is a safer alternative to the intravascular siting of glucose sensors, avoiding the dangers of thromboembolism, as well as of rapid dissemination of infection. Problems of reliable monitoring, though different from those of blood, are nevertheless substantial. The implanted device, through its intrusion into a normal tissue architecture, is perceived by the body for what it is, a disruptive foreign body. As a response, the tissue sets up an intense (acute) inflammatory response designed to degrade, isolate and ultimately reject the foreign material.47 The outcome is locally distorted body fluid composition, i.e., modified functional physiology. As a consequence, no matter how reliable or biocompatible the sensor may be, measurements are thereby performed in an environment that is metabolically distorted and also appears to lose the rapid equilibrium relationship with the local blood and the capillary bed supply (Fig. 5.3). With a low-reactivity material, the acute inflammatory stage subsides to give way to tissue repair, which tends to generate a more vascular environment through capillary proliferation, as well as a matrix rich in fibroblasts. Capillary density in the vicinity of implanted electrodes has been shown to increase with a vascular density maximum at a 50 m distance from the sensor.48 However, measurement distortion was observed (sensor instability and slow response) and was thought to be due to local fluid increase. Whatever the mechanism, the distorted response was only marginally diminished for a tissue-implanted glucose sensor by using slow release of dexamethasone as an anti-inflammatory.49 Due to the presence of soluble bioactive agents, the tissue environment becomes highly hostile to the sensor. Local increases in hydrolase enzyme

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5.3 Three general phases of tissue response at implanted electrodes showing various local reactive phases.

activity (acid phosphatase, alkaline phosphatase, aminopeptidase) have also been observed.50 In addition to enzymic hydrolytic action, the local infiltration of white cells (monocytes and macrophages) augments degradative action, not least through the release of free radicals including peroxide.51,52 The latter is a general part of any non-specific inflammatory tissue response, but has long-term consequences for the integrity of any polymeric component of a glucose sensor. In the intermediate phase of the tissue response, the metabolic activity of inflammatory cells53 leads to a general modification of local tissue metabolic profile, and there are some apparent interspecies differences in this,54 which would make it difficult to extrapolate into man. Much depends also upon the size of the sensor burden upon the local tissue. For example, on one larger scale model system where cellulose sponge was implanted subcutaneously,55 extreme lowering of wound oxygen was observed at five days, as well as sequential variation of pH and CO2. The above changes not only feed back into further tissue reactive responses, but also lead to further development of unpredictable `solution' environments for a sensor. Thus, oxygen and lactate have been suggested as regulators of the wound healing process. High levels of lactate accumulation in wounds may improve wound healing through better collagen deposition, though this is contrary to the requirement for a minimal fibrous tissue in order to achieve undistorted tissue glucose monitoring.56 Quite apart from any physical cell/colloid accumulation, tissue remodelling and local fluid volume change, local lymphatics and capillaries are bioreactive components of tissue. During the wound healing process, they are known to undergo fluctuations in size, fluid throughput and perivascular organisation,

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inevitably impacting on fluid exchanges and solute filtration. These changes may also lead to capillary blood glucose supply variations and, therefore, unpredictable fluctuation in measured levels. Such physiological changes in blood flow were proposed as a reason for fluctuating patterns in oxygen response at a (cerebral) tissue implanted oxygen electrode developed by Clark et al.57 Non-traumatic implantation of a glucose sensor is not feasible currently, but clearly minimal trauma induction is a rational approach, and only truly achievable with a miniature device. Indeed, if dimensions can be made much smaller than those currently available, there would be major gains in reliability. However, as with conventional biomaterials, moving into an era of active control of the local tissue response through a local drug delivery strategy and device-incorporated bioactive agents may also change the outcomes. Thus, it is likely that implantable sensors will, in the future, use bioactive components for conventional implants. This would complement ongoing efforts aimed at the refinement of the bioelectrochemical transduction process itself.

5.4

Materials interfacing strategy

Direct biosample interfacing for a bioelectrochemical device is not feasible. Not only does glucose/O2 permeation into the device require better management in the absence of sample preparation, but also an interfacing material has to be provided to protect the internal components. As a consequence, selective encapsulation is mandatory. Therefore, specific investigation is required of membrane, interfacing and packaging parameters, inclusive of solute (e.g. glucose/oxygen) partitioning and transportation requirements.

5.4.1 Membranes for biosensor interfacing Polymers, rather than inorganic membranes, have been used from the start of biosensor technology to separate sample from the sensing components. As packaging and separation phases, they do offer some useful solutions to the interfacing of biosensors with complex sample matrices. This sub-area amounts to a materials sciences effort in the development of both low fouling and high selectivity barrier structures. Membranes may be classified according to their polymeric constituents (charged, neutral, amphiphilic), and their structural anisotropy. The most convenient and relevant for biosensors is the classification based upon available pore size. On the large scale, there are standard filters of 100 m pore diameter; microfiltration membranes of 1±10 m pore size are used to separate viruses and whole cells; ultrafiltration >10 nm membranes are used to discriminate macromolecules and colloids; reverse osmosis membranes of lower nominal pore size are able to resolve small organics and ions, though in reality these have no discrete pore architecture.

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In the case of porous structures, pore size variation allows for aperture control, and therefore for control over solute access to and from the device. It is likely that commonly used polyurethane external membranes, for implantable (needle) biosensors, operate as glucose diffusion controlling phases through their inclusion of a defined bulk phase porosity;58 certainly the polymer itself is impermeable. As a consequence, membrane technology offers a generic, adaptive platform for the design of a presenting sensor surface. Almost invariably, polymer membranes are used, and variously optimised according to the transport needs of charged, neutral and amphiphilic constituents.59 However, a wide array of chemical and structural features determine eventual interactions, including charge, surface energy, topography and the presence of specific functional groups. Individual examples of improving compatibility include the use of diamond-like carbon (DLC) coatings,60 as non-plasticised, robust, noncrystalline amorphous layers. Here, controlled analyte transport through DLC thickness variation was achieved, with the degree of surface fouling modified through surface profile and surface energy changes. In another example, surfactant incorporation61 into polymeric phases reduced fouling through creation of a quasi-fluid interface with high surface molecular mobility giving a mimic to natural soft material surfaces. The most striking form of biomimicry has involved direct adaptation of natural cell lipid membrane motifs and molecular components. These were embodied in particular in the phosphorylcholine zwitterionic layers pioneered by Nakabayashi and Chapman respectively,62,63 and are modelled on the external surface of the cell plasma membrane. Though these are functionally relatively effective, with phosphorylcholine being the key component, mechanical stability, however, requires covalent attachment to polymeric materials, detracting from the original fluidity properties of the natural plasma membrane. Also, the need for solute (e.g. glucose) permeability limitation through the lipid construct precludes the use of a high phosphorylcholine content. The fabrication of membranes for biosensors encounters various problems, particularly so when it is reproducible and well-defined porosity that is required. The operational problem is due to the presentation of a porous morphology to the sample matrix and thereby due to the tendency to take up protein aggregates into the pores, leading to progressive blockage. Our understanding of the dynamics of the pore blockage in relation to pore size and geometry is developing64,65 but remains incomplete. Progress in this area would be achieved by the rigorous definition of the term `porosity', a parameter unfortunately dependent upon the method used to determine it.66 There is also the possibility of membrane voids permitting antibody/ signalling molecule access to the antigenic constituents in the device, notably the enzyme or any background inert protein (e.g. albumin) to immobilise the enzyme. Moreover, avoidance of any protein release from the sensor demands

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robust chemical immobilisation, but even release of limited quantities of protein if they occur over extended time periods might lead to body sensitisation in susceptible patients. Paradoxically, release of soluble proteins into the body is more likely to be antigenic than that of larger protein aggregations made up of, say, crosslinked material. The membrane interface is thus an important safety feature for a bioelectrochemical sensor. Given the potential dangers of prions and the more classical hazards of viral particles, biocomponent origin, traceability and quality control will be as important to consider as the external membrane barrier in any implantable device.

5.4.2 Membrane property requirements Important surface properties to consider (Table 5.1) depend upon the biomatrix application. Whilst great emphasis was placed in the past on hydrophobicity and a negative charge, in order to nominally reduce the deposition of negatively charged cells and to avoid polar purchase sites for hydrophilic protein domain interactions, in practice, neither approach can be relied upon in most cases. Adverse surface interactions have proven rather more subtle. There is a strong possibility that membrane surface feature size and profile is important, and that an ideally `perfect', smooth surface would encounter minimum biofouling. Table 5.1 Key surface properties of polymeric membranes Chemistry Mobility Topography

Polar, charged, H-bonded, hydrophobic, ionisable Backbone, sidechain, plasticiser mobility Roughness (molecular, cellular), irregular/regular pattern porosity, pinholes

Also surface fluidity on the micron scale, analogous to cell membrane fluidity, may be a means of avoiding protein cell adsorption. Phosphorylcholine (PC), which has both positive and negative charges, is found on the external surface of red cells. It offers an elegant biomimetic solution in order to achieve low fouling.67,68 However, a comprehensive PC layer is likely to inhibit the transport of polar diffusible species and target macromolecules to the sensing element of a biosensor.

5.4.3 Conferred functional advantages of membranes Membrane and thin film technology may well in future make the difference between commercially viable and non-viable devices, viz. biochips. Thus, they confer a variety of benefits, which reduce the high level intrinsic specifications demanded of microfabricated biochips, so serving as a useful complementary

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Table 5.2 Benefits offered to biosensors by use of membrane technologies Membrane technology Surface modification Aperture control Chemical `gate' Sample clean-up Interferent control In-vivo biocompatibility

#

Without surface modifications biochips are vulnerable

technology (Table 5.2). Surface modification allows adaptation of the base materials used for biochip production. Aperture control through pore size management allows for reduction of analyte access to the device and therefore, without sample dilution, enables the device to operate at concentration ranges where the biocomponent binding would normally be saturated. Thus, at high concentrations, bioaffinity is zero order with respect to concentration and, in that case, the biosensor becomes of little use. If the membrane has surface-attached pendent groups, then solute transfer through its pores will be affected by solute/ wall interactions. The greater the interaction, the lower the transport. Such chemical `gating', whilst not often used, could resolve mixtures of similar solutes and so underpin resolution by biochip arrays. Sample clean-up may similarly be achieved. There are significant advantages resulting from allowing biosensor detection reactions to be fully partitioned from the sample itself. For interferent control, a homogeneous or reverse osmosis type membrane is generally used since, in that case, microsolutes need to be rejected.

5.5

Membrane systems used in biosensors

Considerable versatility and ingenuity is required to translate membrane fabrication techniques at the macro-scale to the microfabrication of biosensors. Our past work has emphasised the former and, in general, solvent casting of preformed polymers has been used to coat planar and needle-shaped electrodes. Such techniques could be utilised in order to lay down conventional thick films (Table 5.3) for biochip manufacture, but more attention needs to be paid to edge effects and the ambient solvent evaporation conditions encountered when forming microsurfaces. More amenable to microscale deposition is in-situ monomer electropolymerisation. Whilst a conducting surface is obligatory for such a coating process,

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Surfaces and interfaces for biomaterials Table 5.3 Membrane materials usable for biosensors Membrane types PVC PVC-lipid PVC-surfactant Diamond-like carbon Cellulose acetate Polyurethane

Silane Sulphonated-PEES Polycarbonate Electropolymerised films Bi-layer lipid membranes

deposition can be highly controlled and solvent-related edge effects eliminated. Electropolymerised films, if conducting, can provide a further means of interrogating the biolayer if biolayer binding to a target is associated with charge related or conductivity behaviour modification of the film.

5.5.1 Thick membrane films at electrochemical sensors Microporous membranes Microporous membranes, based on polycarbonate through which linear track micron diameter pores have been etched, are in common use for particle separation. They are an ideal non-swelling structure for partial colloid and cell separation. In the context of biosensors, the pore size (and pore density) may be independently controlled. As a consequence, so is the aperture for driving down transport of a diffusible target solute into the biosensor. For example, in the case of oxidase-based biosensors, Substrate ‡ O2 ! product ‡ H2 O2

5:1

such as those for glucose and lactate monitoring, the low Michaelis constant (Km) is in this way extended beyond the upper concentration range of the fluid to be analysed, O2 demand is reduced and, as a consequence, an effective linear dynamic biosensor range may be engineered. The low permeability furthermore eliminates bulk convective flow influences, so that a sample viscosity and stirring independent system results.69 Microporous polycarbonate may be further modified with an outer coating of diamond-like carbon (an amorphous H, C alloy without a crystalline domain). The diamond-like carbon nanofilm has high haemocompatibility70 and, because it is deposited from a plasma source, exquisite control over coating thickness and therefore membrane aperture can be achieved. Silicones deposited on a microporous membrane71 also allow for controlled pore aperture but, because the silicone is O2 permeable and impermeable to organics, a larger aperture for O2 is afforded than, say, for a metabolite substrate such as glucose, pyruvate or oxalate. This allows oxidase-activated biosensors to be produced with extended linear responses, since the reaction (eqn 5.1) is then

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even less oxygen demanding. Using silane films, variously produced from the crosslinking of di- or tri- halogen substituted silane monomers, also gives a better haemocompatibility than the base material. A planar microelectrode array utilising the oxidase route for measurement could, in principle, be covered with a single microporous membrane and variously coated with modifying films of silicone or diamond-like carbon. This would create a family of sensors with a wide spectrum of sensitivities (calibration slopes) and linear limits that would allow for multiple redundancy, statistical data acquisition, drift free measurement and higher data security. The challenge is for membrane coating techniques to be carried out on a length scale that matches biochip fabrication. Homogeneous membranes for transport control CuprophanÕ is a well-known and well-established regenerated cellulosic membrane used for medical haemodialysis. It is crystalloid-permeable, colloidrejecting, and non-toxic. As such, it has been used in many reports in the past as a covering barrier layer for enzyme-based biosensors. However, its permeability and selectivity cannot be readily altered. On the other hand, PVC is, at first examination, an unpromising material for selective separation or controlled dialytic transport; it is organic solute

5.4 Response to lactate of a lactate oxidase based sensor that has an outer membrane of PVC permeabilised with the surfactant Triton X-100. As the weight-percentage of surfactant decreases the effective linear response and dynamic range of the sensor is extended.

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5.5 The cationic surfactant (MTAC) is incorporated into cellulose acetate (CA) as an outer membrane coating for pyruvate. As the surfactant content increased the effective response of the sensor also increased due to increased retention of co-factor within the proximity of the working electrode.

impermeable. However, it is an excellent support matrix for surfactant.72,73 Figure 5.4 shows the neutral surfactant Triton X-100 physically supported and entrapped in PVC at varying weight ratios and used as an outer barrier membrane for a lactate oxidase enzyme electrode. The lowering of surfactant content reduces an already remarkably high lactate ion permeability to optimise both biosensor sensitivity and the effective Km (linear range) of the integral oxidase of the biosensor. The addition of surfactants to a polymer membrane helps separate polymer chains in order to create voids for increased solute transport. In Fig. 5.5, cellulose acetate was mixed with a methyltrialkyl cationic quaternary ammonium ion surfactant (MTAC), which allowed better transport of anionic pyruvate across the resulting membrane. Importantly, the membrane cation was able to retain positively charged cofactors for the enzyme. As a consequence, only a relatively low drift was observed, even in cofactor-free sample. Porous membranes for lateral transport control The sandwich immunoassay format (antigen capture by antibody, identification via second labelled antibody) is an advantage for electrochemical detection of the binding event. It is necessary to separate attached second antibody from that which remains free in solution. However, in the case of the electrode being an integral part of the flow system, as is shown in Fig. 5.6 (using a system based on a silicon flow channel), the electrode element needs either to be subjacent to the capture antibody, or strategically placed further downstream. Transport of

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5.6 Sandwich immunoassay format within silicon microchannels where the detector electrode is either positioned at or downstream of the immobilised capture antibody site.

reporter antibody in a single step in an open dipstick planar configuration is more difficult, but has the advantage of being pumpless (Fig. 5.7), through exploitation of capillarity. The support material for such transport can be a microporous membrane. However, when nitrocellulose is used, it strongly adsorbs the second antibody (here with an alkaline phosphatase label) and lateral transport is therefore prevented. Nevertheless, the problem can be overcome by preadsorption of a blocking protein (Fig. 5.8). Unfortunately, this strategy cannot be used with cationic nylon for example (Fig. 5.9). As a conclusion, membrane materials have the potential to accommodate fluid microflows and, with judicial choice of the carrier surface (as with surface fouling), the affinity of mobile macromolecules can be modified and their transport behaviour controlled.

5.5.2 Non-conducting electropolymerised films Electropolymerisation of ultrathin films on working electrode materials is a means of film coating that is used with equal facility for macro-scale and microscale electrodes. The basic electrode surface develops a protective layer analogous to that of films used for corrosion protection. One functional aim is to produce low molecular weight cut-off films in order to exhibit high H2O2

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5.7 Open flow planar membrane format in which capture antibody is immobilised on solid phase support; sample and reporter antibody are applied from bulk solution.

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5.8 Alkaline phosphatase labelled antibody transported laterally using albumin preblocking of the carrier surface.

selectivity, which is the primary product for biosensing of oxidase-catalysed reactions (eqn 5.1). Such films may also improve haemocompatibility, extend the linear range for the analyte of interest and help better understand the interactions between sample and biosensor surface. Polyphenolic films: biocompatibility and selectivity The electrooxidation of phenol in solution produces a half wave at about +0.65V vs. Ag/AgCl. No reduction wave can be observed on reversing the voltage sweep in cyclic voltammetry (Fig. 5.10). This is a strong evidence of a stable, inert film on the electrode surface, which accumulates on multiple sweeps until a limiting

5.9 Albumin blocking of nylon does not overcome strong antibody adsorption.

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5.10 Electropolymerisation of phenol (10 mM in PBS pH 7.4) on Pt working electrode. Scan rate 50 mV/s. Following initial oxidative wave at +0.65 V vs. Ag/AgCl due to electrooxidation of monomeric phenol, polymerisation occurs such that Pt becomes insulated preventing further oxidation of phenol. The process is therefore self limiting.

5.11 Effect of polyphenol modification on selectivity profile of Pt.

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insulating layer is formed, which precludes monophenol oxidation. This film layer provides an entirely different selectivity profile for the device (Fig. 5.11), which in particular is seen to be relatively accessible to H2O2, whilst largely eliminating transport of common blood interferents. Catechol, in this case, is not considered as a constituent of blood but as a potential end product of a redox indicator reaction cascade usable for monitoring dehydrogenase reactions.74,75 The haemocompatibility of a surface is usually improved by polyphenol. Although some drift is observed, the comparison with the bare electrode is still highly favourable (Fig. 5.12). These films may also offer some advantages not

5.12 (a) Susceptibility of bare Pt working electrode to whole blood exposure; (b) Stabilisation of Pt working electrode to whole blood exposure through polyphenol modification.

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only for microfabricated structures used in blood but also in protein loaded solutions as they might be considered for proteomic arrays. Polyphenol variants Phenolics with more complex structures and pendent groups would be expected to modify the selectivity and biocompatibility performances. A range of these have been tested (Fig. 5.13). Of these, poly(rosolic acid) appeared to be the most stable (Fig. 5.14). There is further possible manipulation of the interface using different conditions. For example, polymerisation appears to be faster at higher potentials and the net charge requirement becomes lower. These are free radical propagated polymerisations following the initial phenol oxidation therefore forming efficiencies might be expected to be different. Higher solution pH gives more porous films and selectivity is due to a cumulative effect depending on deposition times. Consequently, individual films may be tailored or used in systematically varied biochip arrays.

5.13 Selection of phenolic monomers used during formation o f electropolymerised coatings on Pt working electrodes.

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5.14 Selectivity of Pt modified with poly(rosolic acid) coating following exposure to whole blood.

Polyphenols with surfactant entrapment Polyphenols provide entrapment of `bystander' solutes, in solution, provided that these are of sufficiently high molecular weight. Surfactants are of interest, since not only are they sufficiently large to be physically entrapped in the films, but their solubilising effects and variable charge properties (Table 5.4) would help confer a modified function to the films, with possibly better haemocompatibility. Figure 5.15 suggests that fouling in blood may be reduced over time. In all instances, with film coatings, the level of stability in blood is excellent compared to that obtained with a bare electrode (Fig. 5.16). These data strongly indicate that the base Pt (and other) electrode materials are not sufficiently reliable for use in biofluids, no matter how sophisticated the Table 5.4 Surfactants used for entrapment in polyphenols Surfactant

Polarity

Formula weight

Critical micelle concentration (cmc) (mg/ml)

Non-ionic

8400

100

Adogen 464 (methyltrialkyl (C6±C10) ammonium chloride)

Cationic

368±452

n/a

Taurocholic acid

Anionic

537.7

6.7

Pluronic F68

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5.15 Effects of incorporating surfactant within poly(rosolic acid) film on stability of Pt electrode following exposure to whole blood.

5.16 Reduction in sensitivity of bare Pt following exposure to whole blood, shown as normalised response.

electrode may be (e.g. a biochip array), unless some specific means of protecting against fouling has been used.

5.5.3 Conducting polymer films Planar conducting films, whilst potentially usable as a membrane barrier or as a biocompatibility inducing interface, can also operate as reporters for bioaffinity

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5.17 Species used for the formation of conducting polymer systems.

reactions. Moreover, they can be formed over microelectrodes with the ease of non-conducting phenolics.76 A range of organic compounds have been used to form conducting films (Fig. 5.17), but polypyrrole appeared to be especially advantageous. Conducting polymers can be switched between the conducting and insulating states. Film conductivity correlates with the prevalence of mobile cationic species (polarons/bipolarons) available on the polymer backbone.77 Polaronic conductivity is due to a delocalised positive charge spread over four pyrrole units, whereas electronic conductivity involves the conduction band of the polymer. The films generally used in aqueous solution demand relatively low potentials for redox switching and changes of the state of conductivity. Film deposition is initiated by the oxidation of a monomer to produce a radical cation, which then reacts with other monomers or monomer radicals to go to a dimer stage and then a trimer stage and so on. Oligomer formation then moves on to nucleation and polymer elongation. Importantly, the charged polymer incorporates solution anions. If these anions are bioreceptors, then a biosensor results. The bioreceptor may be co-entrapped in the films without formal charge-based incorporation, provided that a simple anion is present in solution to facilitate the polymerisation process. Inorganic anions as well as anionic oligostructures, such as sodium dodecyl sulphate (SDS) and toluene sulphonate, can be entrapped as stabilising counter ions.

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Conducting polymer electropolymerisation A planar two-electrode arrangement (Fig. 5.18) comprising interdigitated gold (Au) electrodes if used as a polarisable electrode can allow surface deposition of polypyrrole. Once nucleation has occurred, film growth can be followed during cyclic voltammetry (Fig. 5.19). The counterion incorporated affects the voltammetry signatures of the formed films, which is observed when studying the reduction peaks (Fig. 5.20). However, avidin incorporated into a film, when

5.18 Electrode arrangement for planar film impedance spectroscopy. A schematic indicates the 15 micron gap between electrode fingers.

5.19 Formation of polypyrrole as monitored by cyclic voltammetry. There is a continued increase in the magnitude of the oxidative wave with subsequent potential scans.

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5.20 Effect of incorporated counter ion on the reduction peak signature of poly(pyrrole).

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used as a receptor for D-biotin, shows little effect on the cyclic voltammogram. This would seem to indicate that voltammetry is not an appropriate means of detecting receptor changes in the polypyrrole. Some alternative methods of analysis are therefore required. Interdigitated structures (Fig. 5.18) can be bridged by means of polypyrrole. Bridging can then be used to measure film conductivity with any given counterion. To achieve this, a low ionic strength solution is needed to maximise uptake of a receptor molecule and to control the extent of bridging achieved between electrodes. A relatively high polypyrrole concentration is also needed to avoid inhibition of film growth. It is important to recognise that film structure may vary depending on the counterion used. SDS leads to confluent growth and avidin is associated with discoid, multipoint growth. Such variation may have an effect on analyte access to the film, available surface area for binding and the degree to which non-specific surface binding of proteins and cells occurs at the sample interface. Impedance spectroscopy at planar polypyrrole films Once formed on a surface the polypyrrole film can be switched from an insulating to a conducting state based on the impressed polarising voltage. However, a particular redox state can also be sustained, determined by some intermediate polarising voltage. The set redox state can then be followed by twoelectrode electrochemical impedance spectroscopy (EIS). Two electrode EIS can be achieved using a 20 mV RMS sinusoidal potential, and an impedimetric spectrum taken between an extensive frequency range, from 5 Hz to 13 MHz. Independent measures of polaronic and electronic conduction can be obtained, since they take place on quite different voltage oscillatory time scales. A Bode plot is a useful graphical representation of impedance data, combining capacitive current measurement (out of phase with sinusoidal voltage) with impedance measurement (in phase with sinusoidal voltage). Impedance (Z) defines the relationship between applied potential and current and, more 0 formally, is a vector quantity composed of a real (in phase Z ) and an imaginary 00 0 00 Z ‡ iZ where `i' is the (out of phase Z ) part related by the equation Zpˆ complex number and is defined by i ˆ …ÿ1† and the modulus 00 0 jZj ˆ …Z 2 ‡ Z 2 †0:5 . Physical interpretation of the data is problematic. However, the observed data can be modelled using derived equivalent circuits comprising arrangements of resistors, inductors and capacitors. In practical terms, an electroinactive product or outcome of a bioaffinity reaction can be registered and a two-electrode arrangement is usable, provided that the redox state of the film remains stable. Real films show distinct separate zones respectively of electronic and polaronic conduction (Fig. 5.21). In ionic solutions, such as a phosphate buffer solution, polaronic conduction is augmented because the presence of the mobile anions in the film facilitates charged-linked electron transfer.

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5.21 Two electrode impedance spectra at IDEs.

Impedimetric measurements of LH and biotins For a model system incorporating antibody to lutenising hormone (LH), a polaronic phase is identifiable (Fig. 5.22) but, more importantly, there is a capacitance reduction and an impedance increase after binding to LH. This is seen only when a redox cycle is imposed after the binding took place. This provides an indication not only of how antibody can be entrapped within

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5.22 Reagentless binding with LH ± upper capacitance peak, and lower impedance curve trace the zero LH response. LH concentration 100 IU/l.

polypyrrole, while remaining active and accessible to a peptide such as LH, but also of how, in the absence of any apparent voltammetric changes, the impedimetric signature is altered to thereby achieve reagentless binding recognition. The inherent principle should be translatable to other affinity pairs. Thus, for avidin entrapped in polypyrrole, responses are seen for biotin and biotin analogues that are significantly different from the entrapped urease control with regard to phase angle (Fig. 5.23(a)). The biotin derivatives used are amidocaproate succinimide ester (biotin ester) and biotin amidocaproate 3sulpho ester (biotin sulpho). There is firstly an increase in phase angle observed when exposed to phosphate buffer (this is a background effect that would need to be taken into account even if practical measurements are carried out), but the effect of biotin binding is clear-cut. A key feature of all observed binding-induced phase angle changes is the need for one preliminary redox cycling step to unmask the effect on capacitive behaviour (phase angle). Possibly such cycling allows for realignment of polymer chains around the newly formed complex, in turn changing the polymer conformation and, therefore, the ring alignment and polaronic conduction. The importance of the original gold-polypyrrole surface in conditioning the response is illustrated (Fig. 5.23(b)) in the lack of significant biotin response where an electrode had been reused following formic acid cleaning to remove previous films. The subtlety of this effect is underlined by the finding of basic Bode plot profiles for recycled electrodes that are identical to unused electrodes. Thus, either the gold electrode interface is directly important for conditioning cross-IDE impedance, or the gold surface affects the superimposed film structure. The avidin biotin complex is a useful model to study, but the high binding affinity of avidin (KD ˆ 10ÿ15 L/M) does not readily allow concentration

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5.23 Effect of base IDE condition on polypyrrole film formation. (a) unused electrode (b) recycled electrode.

dependent effects to be followed. These, however, can be seen for a LHantibody combination but it is also notable that with the polyclonal antibody preparation used, notwithstanding a high LH affinity, there was little response in contrast to a monoclonal antibody. Possible reasons for this include blocking contaminants in the polyclonal preparation, molecular heterogeneity, distorting polymer conformation changes following binding and antibody aggregates, reducing microenvironmental effects within the bulk of the film structure. The monoclonal-LH system demonstrated no significant response in phosphate buffer, but response to LH was attenuated by pre-incubation of the LH with bovine serum albumin (3 g/L), tested up to 800 IU/L LH. The albumin possibly blocked binding or there may have been bulk solution associated effects with LH that prevented antibody interaction. Whatever the reason, such complex macromoleculer interactions highlight a further challenge to measurements in biological samples. This is in marked contrast to the use of impedimetric measurements at conducting polymers for gas phase affinity sensing, as developed for artificial nose technology,78 where interferences are minimal.

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5.6

Microflows as surrogate, renewable barrier films

5.6.1 Needle biosensors and in-vivo monitoring Notwithstanding the structural elegance and geometrical precision of MEMs based sensors for in-vitro use, in-vivo insertion, even for percutaneous blood sampling and intravenous lines still relies upon needle technology. The fundamental point is that the cylindrical geometry, miniaturisation and high mechanical specification of needles satisfy the dominant issues for healthcare use and take precedence over the structural and length scale elegance of MEMS devices. In the case of silicon-fabricated devices, there is a further concern over long-term implantation, in that dissolution of silicon and local elemental release may be of significance, an issue that needs quantitative consideration if, say, porous silicon is to be used for in-vivo device construction. The needle design has been consistently used for implantable enzyme-based biosensors. Inevitably the greatest attention has been paid to glucose sensing in view of the healthcare importance of diabetes. One classic construction route for this has used two functional layers aside from the chemically crosslinked enzyme layer of glucose oxidase.69 The first layer is a selective low molecular weight cut off barrier based on a sulphonated polyethersulphonepolyethersulphone (SPEES-PES). This robust ionomeric barrier mostly excludes ascorbate, urate and acetaminophen (TylenolÕ) as major electrochemically active interferents on the combined basis of size and charge. The degree of sulphonation is important and may explain why a high selectivity has not been seen by some.79 The outer layer covering the antigenic enzyme layer must interface with the biofluid. It is also a multilayer porous polyurethane serving to reduce glucose transport to the enzyme layer, retain O2 transport (eqn 5.1) whilst providing mechanical integrity and high resistance to the degradative action of free radicals potentially generated during chemical reactions. The laminate formulation used by us is a combination of five pre-polymer polyurethane layers and a single carbonate polyurethane (CorethaneÕ) outermost layer known for its tissue compatibility and low degradation rates.

5.6.2 Open microflow for in-vivo monitoring Needle glucose electrodes used in tissue for short-term (~24 h) monitoring should, in principle, be acceptable for diabetic home use. In practice though, percutaneous, temporary insertion for such minimally invasive monitoring is difficult to achieve. No matter how stable such devices are in vitro, even in blood, the tissue response to the implant, however, demands firstly a stabilisation period of some hours and then in-vivo calibration in the tissue. Indeed, tissue glucose values, both at steady state and under dynamic conditions, show a clinically acceptable match for nearly all current devices. The reasons for this are complex and reflect transcapillary transport lag time, the limit of the

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microcirculation to supply and remove glucose, the nature of the inflammatory exudative tissue response and the distance between the capillary bed and the implanted sensing surface. The last is compounded by the barrier function of the interstitial tissue `mesh'. Open microflow (Fig. 5.24) allows for delivery to the tissue implant site of a film of hydration fluid delivered over the sensing surface of the device and then into the tissue itself. Because subcutaneous tissue mostly operates at a negative hydrostatic pressure, fluid transport does not necessarily require physical pumping and at a net outflow diameter of 0.2 mm, 1±2 L of buffer enters the tissue. The amounts are therefore extremely low, and well within the range of volume exchanges that occur at the capillary bed under physiological conditions, as driven by net positive and negative transcapillary pressures in the Starling mechanism. The externally introduced fluid flow creates a mobile film over the sensing surface that helps to reduce colloid/cell access to the sensor and reduces fouling. Equally important, the juxtaposed tissue is hydrated, and such hydration of connective tissue reduces the diffusional barrier presented to microsolutes. Hydraulic permeability of oedema tissue, for example, is known to be increased several orders of magnitude in comparison with normal connective tissue. The practical consequence of open microflow is that sensor `run in' time is reduced from a typical 3 h down to a mere 30 min. Also the tissue:blood mismatch is reduced as is the commonly observed temporal lag (up to 12 min) for tissue glucose changes. These advantages are most in evidence in the animal model where tissue-related monitoring error appears all but eliminated.

5.24 Percutaneously implanted needle enzyme electrode showing hydration zone at needle tip following pumpless flow of fluid into the dermis.

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Translation into humans shows promise, however, microfluidic design changes are a likely requirement given the interspecies differences in connective tissue composition and capillary bed behaviour.

5.7

Microfluidics and biosensors

While used as `dip in and register' devices, biosensors hold the promise of deskilled analysis. One element of their current attractiveness consists in the fact that they can be produced by microfabrication techniques, and are therefore ready for miniaturisation at single device or industrial scale production. Thus, regardless of this, achieving full analytical benefit depends upon sample presentation. Thus, sample flow, dilution, background composition and the presence of interfering substances all add to the so-called matrix effect in biological samples, and all variously conspire to distort the basic signal readout. It is a fact that sample presentation and delivery have been a much neglected area of research on biosensors. This is also one reason responsible for their lack of widespread uptake, either for extra-laboratory testing, or in drug discovery arrays. However, thanks to rapid advances in microelectromechanical systems (MEMs), it is now possible to design and manufacture high precision flow structures,80 predominantly in silicon or glass. By using these structures, researchers have been able to create designer fluid flows that offer the prospects of precision fluid delivery to targeted locations via valve controllable flow paths. Beyond any size scaling effect, however, fluids under microfluidic flow conditions exhibit quite distinct fluid dynamics. Here, viscous forces are dominant, and any irregularity or turbulence in the flow pattern evens out in order to give a laminar flow, with precise, quantifiable solution transport profiles through an entire length of the microchannel network regardless of channel interconnections and overall geometry. Microchannels or capillaries (usually 0.02 to 2 mm diameter) can be regarded as reaction microenvironments. They exhibit two very specific characteristics. First, they present small internal volumes that can be easily and rapidly filled by fluids. Second, fluids moving inside these channels remain laminar even at very low velocities, with no evidence of turbulence. A microchannel thus constitutes a special delivery vehicle for self-contained and finely controllable sample exposure of a strategically positioned biosensing surface along the flow path. Moreover, the use of miniaturised flow cells giving Reynolds numbers of less than 2000 (vide infra) permits non-turbulent fluid flow so independent, parallel flow streams can be accommodated without turbulent mixing.81 A microfluidic channel presents a relatively high surface area to the bulk sample volume, which thereby enables a wall-immobilised reagent to be highly effective in reacting with bulk solute constituents, e.g., through binding and catalysis. These high surface to bulk interactions also facilitate chromatographic principles of sample separation.

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Laminar flow allows delivery of chemical species inside capillaries with remarkable precision. The application of different streams of a multiphase, laminar flow system also allows a reaction to take place exclusively at certain regions of the flow system, either between a stream and channel wall or between two adjacent streams. As a consequence, such devices could be potentially used to carry out in-situ microfabrication experiments, thus paving the way to innovative micro and even nanotechnology strategies.

5.7.1 The Reynolds number In 1883, Osborne Reynolds demonstrated that, under specified conditions, stable laminar flow switched to unstable turbulent mode. He introduced a fundamental equation, which gave a range of dimensionless numbers, the Reynolds numbers, defining the transition from laminar to turbulent flow. The Reynolds number is defined as the ratio of inertial vs. viscous forces and is used in momentum, heat, and mass transfer to account for dynamic similarity It is calculated using: Re ˆ

D
V
 

5:2

where Re ˆ Reynolds number, D ˆ characteristic length, V ˆ velocity,  ˆ density and  ˆ viscosity. Due to the large number of parameters, there is a great degree of flexibility available regarding the flow cell construction and arrangement. Typically, a single flow within a straight pipe with smooth walls and Reynolds numbers below 2000 is both stable and laminar. Between 2000 and 2300, a transitional flow is observed. Above Reynolds numbers of 2300, the flow becomes turbulent and unstable. Most microfabricated fluid systems operate under low Reynolds number conditions. Since in small channels only small liquid masses are moved and flow velocities are typically low, this ratio is generally small for any given viscosity. Under these conditions, an aqueous stream in a microchannel will behave similarly to viscous oil-like flow at the macroscopic level. Consequently, it is possible to flow two aqueous streams of similar viscosity side by side while preventing turbulent mixing.82 Nonetheless, if flow is subject to an abrupt change of channel geometry or encounters obstacles, it will change from a laminar to a turbulent regimen even at extremely low Reynolds numbers (Re40-300).81 Although turbulent mixing does not normally occur at low Reynolds numbers, diffusion still occurs, and has been utilised by Manz and co-workers to generate microfluidic mixing, otherwise difficult to obtain at such low Reynolds numbers.83 A key consideration when using microfluidics is to understand and control flow characteristics (velocity profile, shear force, etc.) which is a complex area of fluid dynamics that relies heavily on computational modelling in order to

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predict these outcomes. The Navier-Stokes equation is often used to characterise fluid flows provided that the initial conditions are properly established.84 The equation is based on the balance of forces summarised as follows: Transient Convective Body Pressure ‡ ˆ ‡ inertial force inertial force force force

‡

Viscous force

5.3

However, in steady flow, the transient inertial force vanishes. Also, if the fluid is considered as ideal, the viscous force term disappears, simplifying analysis.85

5.7.2 Design and fabrication of microfluidic devices The main trend currently is to consider the biosensor as an integral part of a complete miniaturised analytical system rather than as a separate entity, thus the concept of a micro total analytical systems (TAS) involved microfluidics as part of the system.86,87 A scaled down fluid delivery system can be seen as equivalent to the distributed capillary networks of natural tissues, which permits conservation of sample volume and exquisite flow control and fast switching of flow paths and delay lines. All contribute to the desired, near-ideal conditions needed for the fine-tuning, multiplexing and efficient operation of bio- and other sensor systems. Microfluidics is essentially a basic, though less discussed, component of a TAS. Here, the manipulative steps of an analytical laboratory, notably sample separation, clean-up, reagent introduction, reaction profiling concentration optimisation and analytical endpoint detection are incorporated into a miniaturised structure, alternatively considered lab-on-a-chip. Miniaturisation of many of the relevant processing components has certainly been achieved, but none would be operational without the mobile element embodied in the integral microfluidics; much more needs to be done in developing the technology. Microfabrication of biosensors scales down reagents requirements, a major financial advantage, and also facilitates greater sample throughput. The method requirement is for microfluidics. Before fabricating a microfluidic system, time has to be spent in considering the precise application. Thus, the eventual operating conditions need to be specified. One important choice is that of the basic material (resin, polymer, metal, glass or silicon) from which the device will be fabricated. The material has to be chemically inert and resistant to the various solvents and solutes to which it is exposed during the operational lifetime. Each type of material is associated with different types of manufacturing methods and quite different manufacturing costs, which can be more or less important depending on the individual material used. When an aqueous phase is used, relatively inexpensive polymeric materials like poly(methyl methacrylate) (PMMA) and poly(dimethylsiloxane) (PDMS) can be used to fabricate flow-cells. However, some solvents will alter if not

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completely erode these polymeric materials. As a consequence, for a microfluidic device handling a wide variety of solvents (aqueous to organic, polar to apolar), it becomes vital to choose more expensive materials that are chemically inert vis-aÁ-vis these solvents, as well as the solutes they will be carrying. Such materials (metal, glass, silicon) require more complicated, and therefore more expensive, techniques for etching and assembly, reversibly where possible, into a well-sealed microfluidic enclosure. It is also crucial in this instance to choose the right component to dispense and transport the fluids from a reservoir into the microfluidic device. Syringes have to be of glass and the tubing and connectors made of solvent-resistant but still flexible polymeric material, like PTFE (polytetrafluoroethylene) or PEEK (polyetheretherketone), as commonly used in liquid chromatography. Once the choice of material has been made, various etching/assembly techniques available need to be investigated. Microfabrication techniques, in their most basic form, rely on thin-film or thick-film metal pattern formation associated with photolithography.88 This technique has already proved highly efficient and successful in the mass production of silicon chips for the semiconductor industry. A thin metallic layer is deposited onto a substrate (silicon or glass), typically via spin-coating, after which a photoresist is added. Then, a patterned photomask is applied and the whole system irradiated by UV light. The unmasked portions of the photoresist are etched and the exposed metallic areas are then eroded either chemically in solution (wet etching) or by exposure to a plasma (reactive ion etching).89 The metal layer can be replaced by a layer of silicon oxide (SiO2) that is grown on a pure silicon substrate in a furnace over many days. The SiO2 layer can be etched using the same techniques as above. Silicon oxide and glass, which are similar, can be easily etched using HF (wet etching). The microfluidic channels formed can be sealed by anodic bonding of a glass slide or cover slip to the glass or silicon/silicon oxide substrate. The lift-off technique is very similar to the classical photolithography technique except that the photoresist is first deposited onto the substrate, then partially etched by UV light using the patterned photomask. The metal layer is then deposited on both the exposed substrate layer and the remaining photoresist. The photoresist is then dissolved (lift-off) leaving the patterned metallic layer. Anisotropic etching is capable of producing three-dimensional structures as opposed to photolithographic techniques. It is performed with heated alkaline solution using SiO2 or Si3N4 masks. The etching rate is highly controllable and stops at layers of impurities. It is dependent on the crystalline orientation of the substrate. In silicon etching, for example, the most commonly used etching agents are potassium hydroxide and tetramethylammonium.90 Powder blasting is a technique that can be used to generate microstructures in brittle substrates such as glass. It employs a masked substrate and an eroding

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beam of 30 micron sized alumina (or sand) particles in a powder form. This powder is propelled using a compressed air stream, which is able to etch channels down to 100 micron width and 20 micron depth.91,92 Many of the techniques employed for microfabrication are borrowed from the semiconductor industry. However, since silicon is hydrophobic by nature, it may not be suitable for biological studies, and the use of polymers represents a more attractive and cheaper proposition. Poly(methyl methacrylate) (PMMA), for example, can be used and is easily etched using standard mechanical tools or using laser ablation techniques to form very precise microchannels. Hot embossing is a technique capable of generating low-cost devices for the biotechnology industry. It utilises a hot plate with the substrate being pressed between the hot plate and another patterned metallic plate. Micro-injection moulding requires that the polymers used are first reduced to granules, melted above their Tg (glass transition temperature) and injected into a mould, generating microfluidic channels. Another polymer that can be used is Poly(dimethyl siloxane) (PDMS), a clear rubbery, crosslinked polymer that can be easily cast into a patterned metallic or silicon mould, thus generating microfluidic devices at very low cost, ideal for prototyping.93,94

5.7.3 Detection Chemical and biosensing devices integrated into fluidic arrays at multiple locations can allow sequential monitoring of reaction transients or of the kinetic development of reaction profiles along a given flow channel. Little additional structural distortion of precision-formed fluid channels results with wall embedded devices. More accurate kinetic measurements thereby achieved thus can add to end point estimations. In electrochemical sensors oxidisable or reducible molecules generate measurable current flow. With stable fluidics these current responses become further stabilised, and signal to noise is reduced, important for low current outputs. Neurotransmitters and the products of some enzyme reactions (oxidoreductases), either directly or via electron transporting indicators, have been measured in this way and even when using well-known chemistries there are additional benefits due to miniaturisation and the fine tuning of signal outputs. Through a combination of biosensor arrays and variable flow conditions, a spectrum of multiple responses can be obtained giving better sample profiling allowing for greater biosensor redundancy and ultimately improved data reliability. Sample concentration changes are also possible in miniaturised structures through controlled evaporative loss and, in the case of ionic solutes, adsorption/release from electrically addressable surfaces give additional routes to sample adjustment thus allowing the analytical `reach' of biosensors to be extended.

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Control over diffusive solute transport from the bulk sample to the biosensor surface can be augmented by in-situ polymerisation of thin conducting and nonconducting films (vide supra), readily formed within the confines of a closed microchannel. Such thin films do not distort flow, but do allow co-entrapment of biological or non-biological reagents for biosensor formation. Fluid conductivity measured with micro-electrode pairs across flow channels or specific ions measured with miniature ion-selective electrodes allow for both profiling of a sample and facilitate sample status monitoring in addition to any integrated biosensor for a specific target. A special opportunity arises through optical fibre or waveguide based optical interrogation of flow components, and detection of an intrinsic solute property (e.g. protein fluorescence) opens up the additional armoury of spectroscopy.

5.7.4 Diffusion and sample separation Laminar flow in microchannels, depending on the flow rate, may allow or prevent the diffusion of solutes from one stream to another. Diffusion becomes more prevalent as the flow rate decreases and the flow path increases. It is also possible to discriminate between analytes depending on their molecular weight. Low molecular weight compounds diffuse more readily than high molecular weight solutes for which diffusion can be almost completely prevented, reducing fouling of a sensing electrode surface. Figure 5.25 shows a schematic microfluidic device, where a bifurcation is used to converge fluid streams in a 200 m wide flow channel. The value of this flow phenomenon is that there is controlled laminar flow over any electrode located in the wall of the flow channel, so convective transport and therefore response of an electroactive species can be correctly modelled. Furthermore, a protective fluid film can be used over the intramural electrode to compartmentalise the electrochemical (or chemical) detection sequence from a potentially contaminating

5.25 Bifurcated flow cell schematic for non-mixing of parallel microflow.

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sample stream. The protective effect primarily applies to microsolute monitoring, where diffusion coefficients are at least one order of magnitude greater than for, say, surface-active proteins; cellular transport, moreover, across the two streams is virtually eliminated. Even with a protruding electrode it may be possible to have a viable monitoring set-up with the protruding device allowing for a high exposed microsurface area and therefore higher currents. In a study on the difference of diffusivity between a high molecular weight solute (glucose oxidase) and a small molecule hydrogen peroxide (H2O2), generated by the reaction between glucose and glucose oxidase, it was shown that with glucose oxidase in the stream adjacent to the electrode and glucose in the opposite stream, a reaction occurred at the interface between the two streams, generating H2O2. With increasing flow rate, the electrochemical current generated by hydrogen peroxide at the electrode decreased and a faster flow reduced the diffusion time for hydrogen peroxide. This behaviour is similar to that of a separation or filtration membrane but in a fluidic, thus renewable, form preventing biofouling of the sensitive sensing surface.

5.7.5 In-situ microfabrication Many experiments have been carried out using the unique behaviour of fluids inside microfluidic channels. Control over reagent delivery has been used to precipitate solid structures in flow channels such as linear wire electrodes or silver and conducting polymeric fibres.81,95 However, microfabrication of complete barrier structures, like free-standing ultra-thin membranes remains a challenging task. Thus, attempts at precipitating cellulose acetate from 0.2 wt% acetone solution at an aqueous stream using first a 1.5 mm (width)  0.5 mm (depth) channel and then a 1.2 mm (width)  0.6 mm (depth) channel led to aggregates of precipitated polymer but not membranes. The formation of distinct membranes with some mechanical strength was achieved by using a polymer that was not easily redissolved by the flow stream. Thus, interfacial polymerisation of nylon 6,6 was attempted. Adipoyl chloride (62.5 mM) in xylenes and 1,6-diaminohexane (62.5 mM) in distilled water were used. Interfacial polymerisation in the diffusion zone led to nylon 6,6 membrane instantly. A solid anchoring point proved to be a key requirement in order to achieve stable membrane attachment. Further development of permeable membrane structures should allow for an interfacial barrier for sensor protection, and could complement separation by liquid-liquid interfaces.

5.8

Conclusion

Though biochip-based and other precisely configured biosensors are precision structures with the opportunity of greater measurement reliability using multiple arrays, the need remains for reliable and stable interfacing. Relevant research in the

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areas of materials, membranes, fluidics and films is already available to help adapt and `refashion' biochip surfaces for practical use. A combined effort at this stage is more likely to help deliver viable products for the future, especially in healthcare. Admittedly the advantages may be of overwhelming importance only if measurement is required on a continuous basis or if unmodified (e.g. undiluted, non-deproteinised) samples are to be assayed. However, the over-the-counter medical diagnostic market is a prime example where unmodified samples will be measured, which is likely to benefit from biochip configurations of biosensors as truly reliable and therefore `smart' systems, provided that stability in the sample can be guaranteed. To this end, increasing interest in barrier membranes is anticipated. However, fluid interfaces and in-situ microdeposited membrane barriers could lead to major augmentation of microdevice performance with considerable economy of sample volume requirements.

5.9

Acknowledgements

The authors would like to thank EPSRC, JDFI, BDA and the EU for funding of the ongoing studies presented in this review.

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6 Micro- and nanoscale surface patterning techniques for localising biomolecules and cells: the essence of nanobiotechnology Z A D E M O V I C and P K I N G S H O T T , The Danish Polymer Centre, Denmark

6.1

Introduction

Most reactions in biology occur not in solution, but at interfaces, such as cell membranes. The need to improve the quality of our lives has led to substantial research and development in manipulating material surfaces in an attempt to mimic and control biological reactions that take place at interfaces. Medical implant materials, for example, have played a pivotal role in bringing surface concepts to biology. In the late 1940s the first biomaterials, as we known them today, were developed.1 Almost in parallel, scientists began studying the properties of biomaterial surfaces, which included investigations involving protein-surface and cell-surface interactions, and ways of modifying these surfaces in order to control so-called biointeractions. Present surface science has had considerable impact on biology and medicine. Implant biomaterials,2 affinity chromatography,3 gene chips4 and diagnostic arrays,5 cell culture surfaces6 and biosensors7 are examples where surface technologies have been applied to biological problems. These applications have been largely driven by an early appreciation of the fact that surfaces provide the platform for biological reactions to take place. An optimally functioning human being is made up of tissues that consist of highly organised structures of different cell types, each performing different but complementary functions. All the cells are surrounded by topographic and chemical cues that can align and guide cells.8,9,10 The specific protein-surface and cell-surface interactions are most likely to be strongly influenced by a combination of surface chemistry and surface topography,11 arranged, moreover, into well defined patterns on the surface at both lm and nm length scales.12 The relative importance of these two factors remains poorly understood because we are still developing techniques to create patterns and the analytical methods capable of actually visualising such features. In addition, results achieved so far have shown that topographical modifications are accompanied by chemical variation. Hence, controlling the topography and chemistry of polymer surfaces independently is an important issue. The effect of surface topography on cellular

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and tissue response to implanted biomaterials has been known for the past 20 years. Cells have been shown to respond to micro- and nanoscale structures and to topography. Slightly roughened surfaces show increased osteointegration, reduced fibrous encapsulation and enhanced integration of implant materials,13,14,15 a result of increased adhesion of connective tissue cells. An ability to properly engineer cell-surface interactions and to control the spatial location of cells is highly desirable, and essential for creating in-vitro mimetics of physiologically relevant interfaces. Understanding how the surface chemistry and structure of a material can be used to control the biological reactivity of a cell interacting with that surface is the ultimate challenge for biological surface science. Ideally, the surfaces will need to have well-defined arrays of biorecognition sites designed to react specifically with the cells, since many of the important functions of cells depend on the arrangement of biomolecules at their surfaces.16 A large number of fabrication tools exist for tailoring a biomaterial surface. Chemical surface modification can be combined with designed microstructures and nanostructures, aimed at matching biological components or for inducing a desired biological reaction. Controlled protein adsorption is crucial to optimise cellsurface interactions, since cells recognise the protein layer and not the underlying surface. Surfaces must be developed that control the conformation and orientation of proteins with precision so that the body will specifically recognise them. The simplest method for immobilising a biomolecule on a surface is physical adsorption, but this is a purely random event and can lead to denaturation and loss of protein function or displacement by less desirable proteins. A more stable kind of protein immobilisation involves creating covalent bonds between the protein and solid surface. In addition to efforts made to optimise and control protein orientation at surfaces, a new research direction in protein immobilisation is taking place, namely protein patterning.17 It can be defined as protein immobilisation within specific locations in either two- or three-dimensional space.18,19,20 To date, research in this field has had considerable success on the micrometre length scale with patterning in two dimensions, primarily with single proteins. The generation of small structures is central to modern science and technology. The most evident examples are in microelectronics, where smaller has meant less expensive, faster, more components per chip, higher performance and lower power needs.21,22 Micro-fabrication techniques, which have been developed for application in the electronics industry and in information technology, have been progressively employed in biomaterials research to study cell-surface interactions.23 One aim of these studies is the spatial control of cell attachment and organisation.24,25 The ability to generate patterns of proteins and cells on surfaces is important for biosensor technology,26 tissue engineering,27 and fundamental studies of cell biology,28 and can generally be called the field of nanobiotechnology. The positioning of biological ligands at well-defined locations

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on substrates is required for biological assays,29 for combinatorial screening,30 and for the fabrication of biosensors.31 Tissue engineering requires cells that are placed in specific locations to create organised structures in three dimensions. Furthermore, the growing demand in biosensor technology for high-density, high-sensitivity, multi-analyte chips can be met only with precise and reproducible patterning technologies that allow controlled positioning of chemically different active areas. The similarity of needs and limits between the implant and biosensor fields has led to the development of chemical, topochemical and topographical patterning methodologies that are applicable to both areas. Strategies that have been explored for fabricating patterned nanostructures take advantage of at least one of the following principles: · interaction of matter (`lithography') with photons (x-ray,32 ultraviolet33), energetic particles (electrons, ions, neutral metastable atoms) and scanning probes34 · replication against masters via physical contact35 · self-assembly of molecules36 and nanoscale objects.37 Most of these methods involve sophisticated instrumentation or treatments, and only a few of them are suited to create topographical variations at polymer surfaces without changing the chemistry. Photolithography is the most widely applied technology in the fabrication of microelectronic structures but this technique is expensive and too slow for the production of nanometre features. Self-assembly techniques provide the opportunity to produce features with a higher resolution down to a few nanometres. Micro-contact printing allows fast and reproducible production of chemical micro- and nanostructures with a resolution as good as 50 nm. Smaller dimensions are limited by the contact stability of the stamp with the substrate. Dip-pen nanolithography, blockcopolymer lithography, colloidal lithography, and self-assembled monolayer lithography are additional techniques that use self-assembly strategies to produce nanoscale structures. These methods are described in more detail below.

6.2

Lithographic patterning with photons, particles and scanning probes

Photolithography (Fig. 6.1(a)) and electron-beam (e-beam) lithography (Fig. 6.1(b)) are the most frequently used techniques in micro- and nanotechnology, used to manufacture components for computer technology. However, they have also been used most extensively to create features for patterning proteins and cells. Photolithography uses electromagnetic radiation to introduce a latent image into the appropriate material.38,39 The surface is first covered with a radiation sensitive film, usually a polymer called `resist', then exposed to a beam of radiation, which modifies polymer properties at the irradiated areas. Subsequently, a patterned structure can be developed through etching or

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6.1 (a) Schematic of the general principle of photolithography and patterning with self-assembled monolayers (SAMs) (e.g. on Au). A photoresist is deposited on the substrate (e.g. by spin-coating) of choice and light is shone through the mask of defined pattern creating a pattern. A SAM with one type of head group (thiol 1) is chemisorbed onto the exposed Au substrate. The leftover photoresist is then dissolved away and a second SAM (with a different head group) is chemisorbed into the gaps, thus creating a surface with different chemical patterns. (b) Schematic of lithography using either electrons (ebeam) or photons. The principle is the same as (a) except the pattern is written directly into the photoresist followed by developing with a solvent. In e-beam lithography features down to 5±10 nm are easily and reproducibly fabricated.

dissolution of the image, leaving a pattern of polymer on the surface that serves as a mask for further surface treatment of the uncoated areas. Different variations of these techniques, such as the use of short wavelength light sources, for example deep UV or X-ray, the chemical adjustment of the polymer resist material to the light source has allowed further decrease of patterned dimensions. The major drawbacks with conventional lithography are the large costs due to cumbersome fabrication processes and low throughput rates for nanometre sized features. In addition, photolithography is also not well suited for the modification of surfaces involving delicate biomolecules.

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6.2.1 Photolithography (smallest feature size obtainable is 50 nm) Photolithographic micro-fabrication of surfaces is a well-established technique for the production of model surfaces with defined topographies or with chemically defined patterns. Photolithography with photons uses a mask or a hole to localise the beam spatially. This technique is one of the most widely used in science and technology. The resolution of photolithography increases as the wavelength of the light used for exposure decreases. On the other hand, as structures become increasingly small, they also become increasingly difficult and expensive to produce. With 193 nm light from an ArF excimer laser, patterning of features down to 150 nm has been achieved.40 Patterning of features below 100 nm becomes challenging because of the lack of transparent materials suitable for lenses or for support for photomasks at wavelengths below 193 nm. Sun et al. developed a scanning near-field photolithography (SNP) technique to pattern self-assembled monolayers (SAMs) on gold with feature sizes of 40±50 nm (Fig. 6.2). These materials have been used as masks for pattern transfer to the underlying gold substrate by wet chemical etching creating three-dimensional architectures.41 Chemical modification of the terminal reactive group of patterned SAMs will enable development of materials of interest and find applications in diverse areas of nanotechnology.42 Photolithography and focused laser methods have been used to pattern surfaces with molecular layers,43,44 but photolithography requires the use of harsh solvents and bases, making it incompatible with many biological molecules. The laser method uses an interference technique that does not allow generation of patterns of arbitrary complexity. However, protein patterning using chemical linkers to create a heterogeneous monolayer is achievable.44 Silanes have been a widely used reagent to attach proteins to silica or metal surfaces because they can withstand the harsh solvent system required to remove the resist. Textor et al. describes a photolithographic technique based on evaporated metal films of titanium, aluminium, niobium and vanadium with patterns in the 50±150 lm size range to study protein and osteoblast cell interactions with different geometries and chemical composition of the patterns.45,46 Spatz et al. used ArF excimer laser ablation to generate selforganised, grating-like periodic structures on the surface of PET films.47 They demonstrated that both the morphology and the orientation of the melanocytes are determined by the topography of the PET substrates structured by the laser. A novel PEG surface modification and patterning process has been proposed.48 This approach mimics traditional photolithography in that spin coating and UV exposure through a photomask are employed to create a polymer pattern. The obtained PEG hydrogel microstructures resisted protein adsorption and cell adhesion whereas non-modified substrate surface promoted cell adhesion and spreading.49

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6.2 Generating nanostructures by scanning near-field photolithography (SNP). In this example a combination of self-assembled monolayers and wet chemical etching are used to create the patterns. (a) An AFM topography image reveals a set of 55  5 nm parallel trenches etched into a gold film. The SAM used was a HSC15CH3 molecule chemisorbed to the gold surface. The wet chemical etching was performed using a solution of Fe(CN)62+/Fe(CN)63+. (b) The cross-sectional topography trace orthogonal to the lines (taken from ref. 41, with permission).

Other problems accompanied with photolithography are damage and deformation of the mask on exposure to energetic radiation and the high cost and low speed of this technique. Furthermore, it cannot be easily applied to nonplanar surfaces, tolerates little variation in the materials that can be used and provides almost no control over the chemistry of the patterned surface. Plasma lithography Goessl et al. have developed a method for the patterned immobilisation of cell surface receptor ligands on biomaterials surface based on the combination of radio-frequency glow discharge plasma (RFGD) and photo-microlithography.50 The method called plasma lithography combines the high spatial resolution of photolithography and the chemical versatility of plasma surface modification. In the first step a non-fouling background is deposited onto the polymer surface, and the surface is then coated with a positive photoresist that is exposed to UV

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through a photomask that defines the pattern. In the second step, plasma treatment, a fluorocarbon plasma polymer is deposited onto this patterned surface. Surfaces composed of micro-patterned domains of fluorocarbon polymer can control the shape and size of vascular smooth muscle cells (SMC). Spreading restricted cells formed a well-ordered actin skeleton, and the cells were still confined to the areas of the adhesive pattern after two weeks. Dai et al. generated high-resolution surface patterns of various surface functionalities through the H2O-plasma etching using a TEM grid as a mask.51 They also extended this technique to include conducting polymers using the plasma-patterned, metal-sputtered substrates as the electrodes for electropolymerisation. The regions of the electrode surface covered by the plasma polymer are electrically insulating and hence inactive toward electropolymerisation, whereas the uncovered areas can effectively initiate electropolymerisation. Thissen et al. described a controlled excimer laser ablation technique of a plasma polymer-PEG modified substrate to fabricate surface patterns suitable for spatial control of protein adsorption and subsequent cell attachment.52 By using appropriate masks in the laser beam, a resolution of approximately 1 m was obtained.

6.2.2 Electron beam lithography (e-beam) (feature size obtainable down to 10 nm) In e-beam lithography, a focused electron beam is used to form patterned nanostructures in an electron sensitive resist film.53 Interaction of the electron beam with the resist causes changes in solubility and the formation of local spots that become soluble in a developer. Pattern formation using electrons is a method capable of forming patterns with nanometre resolution in a resist film. With e-beam writing resolution of 10 nm was achieved more then 20 years ago,54 conventional lithography with focused beams of electrons is slow and expensive.55 Consequently, e-beam lithography is more suitable for producing photomasks for optical lithography. Plastic surfaces have been patterned by combining homogeneous polymer grafting with e-beam irradiation and localised laser ablation of the grafted polymer,56 thus fibronectin was adsorbed selectively onto ablated domains and hepatocytes adhered specifically onto the ablated domains adsorbed with fibronectin.

6.2.3 Scanning probe lithography (SPL) (feature size obtainable down to 1 nm) Scanning probe lithography uses a small tip scanned near the surface of a sample to image and modify surfaces with atomic resolution.57,58,59 Advantages of SPL are the ability to generate features with any geometry as well as patterning of

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non-planar surfaces.60 Additional advantages of such an approach include simplicity (no need for clean room fabrication environment), the ability to pattern proteins under non-denaturing solution environments, and direct interrogation of protein binding events. Recently, SPL methods have been explored for patterning protein surfaces. Examples include the mechanical scraping of open patches within protein-resistant polymer films for subsequent protein adsorption,61 application of scanning electrochemical microscopy to derivatise electrode surfaces,62 and the use of nano-grafting to incorporate reactive sites into self-assembled monolayers.63 This method offers the possibility of creating as well as of reading extremely high-density protein arrays.64 SPL is used to scratch nanostructure in soft materials,65 to change the head group or packing density of organic monolayer catalytically66 or to write 30 nm patterns of alkanethiols on gold.67 SPL is a slow method because the writing speed is limited by the mechanical resonance of the tip and piezoelectric elements that maintain the constant separation between a tip and sample. For that reason SPL is better suited for the formation of masters than for replication, but efforts to develop faster SPL are continuing. The use of multiple-tip systems has increased production speeds remarkably.68

6.2.4 Dip-pen nanolithography (DPN) (feature size obtainable less than 100 nm) In dip-pen nanolithography (DPN), an atomic force microscope (AFM) tip is inked with a material known to self-assemble on a solid substrate, as depicted in Fig. 6.3. The tip is brought into contact with the surface, a water meniscus is formed and adsorbed ink is transferred to the substrate when the probe is held in contact or moved along the surface.69,70 The first DNP investigations used thiolbased ink and gold substrates67 but later it has been extended to other inks and substrates. Various oligonucleotides functionalised with a hexanethiol linker have been inked on gold.71 Collagen was also written on a gold surface by cysteine binding on the gold without destroying the helical structure of the collagen.72 Thiolated collagen and collagen-like peptide has been patterned on gold with line widths as small as 30±50 nm, the largest molecule thus far positively printed on a surface at such small length scales. Moreover, DPN can be used to generate complex multi-component nano-arrays of native proteins that are biologically active and capable of recognising a biological complement in solution.73,74 Other inks that do not chemisorb to the surface are being developed. Fluoroscently labelled proteins can be written on modified glass surfaces.75 Utilising electrostatic interaction between positively charged parts of a protein and negatively charged silica, proteins have been directly written onto silica surfaces.76,77 Hyun et al. developed a method to fabricate chemically reactive

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6.3 The concept of dip-pen nanolithography (DPN). Schematic of the process using an AFM tip to deposit chemical moieties on surfaces after it is wetted with the desired molecule of choice. (Taken from Ref 67, with permission).

nanoscale features by patterning a SAM of a COOH-terminated alkanethiol on a gold substrate by DPN, followed by covalent immobilisation of biotin on the nano-patterned SAM and subsequent molecular recognition of streptavidin from solution.78 The resulting streptavidin nano-pattern provides a universal platform for molecular recognition-mediated protein immobilisation because of the ubiquity of biotin-tagged molecules. By this method, periodic arrays of biotinBSA with feature size of 230 nm were readily fabricated.

6.3

Soft lithographic techniques

Current lithography techniques appear to have reached their limits in terms of minimal feature size and fabrication cost. Thus, alternative methods need to be explored to go beyond these limits. An attractive candidate for achieving nanoscale structures is the phenomenon of self-organisation, in which certain materials under certain conditions will arrange themselves or self-assemble into stable, well-defined structures, e.g., when deposited onto a surface.79,80 The development of self-assembled systems has been a major advance in material fabrication technology during the last ten years.81,82 Self-assembly concepts originate from biological processes such as the folding of proteins, the formation of the DNA double helix and the formation of cell membranes from phospholipids. In self-assembly, subunits spontaneously organise into stable, well-defined structures based on non-covalent associations. Self-assembled monolayers

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(SAMs) provide well-defined ordered structures and chemistries that can be systematically varied. Also, spatially defined arrays of SAMs can be prepared by combining self-assembly with patterning methods such as micro-contact printing83 and photolithography.84 Self-assembly is being examined for patterning at scales greater than 1 m and applications are moving to smaller dimensions. Examples of self-assembly as a method for fabricating nanostructures include micro-contact printing of self-assembled monolayers,85 reactive ion etching with thin films of block copolymers as a mask86 and the synthesis of mesoporous materials with aggregates of surfactants as templates.87 Typically, this is a non-serial process (since the whole system organises itself at the same time) allowing patterning of large areas with high short-range and some long-range order. The thickness of the SAM is usually 2±3 nm, and can be tuned with a accuracy of 0.1 nm by varying the number of carbon atoms in say the monolayer alkyl chain. Polymers and block copolymers are an interesting class of molecules for nanomaterials fabrication88 and show a broad variety of self-organising patterns; such polymeric nanostructures find applications in diverse areas that include optics, biochemistry, and material science. The ability to use different molecules with well-defined chemical end groups adds to the flexibility of this approach to create nanoscale patterns of specific functionalities. Soft lithography has been developed as an alternative to photolithography and comprises a number of lithographic techniques such as micro-contact printing (CP),89 replica moulding (REM),90 micro-transfer moulding (TM),91,92 micromoulding in capillaries (MIMIC),93 and solvent-assisted micro-moulding (SAMIM).94 In soft lithography, an elastomer is cast against a rigid master (silicon wafer) and the elastomeric replica subsequently used as the stamp, giving structures that can be as small as 100 nm. Masters are made using conventional high-resolution nano-lithographic techniques. In this way it is possible to produce multiple copies of indistinguishable nanostructures from a single master, rapidly and economically.95 The common feature of these techniques is the use of a patterned elastomer (usually poly(dimethylsiloxane (PMDS)) as the mould, stamp or mask to generate or transfer a pattern to the substrate. In addition, soft lithography uses flexible organic molecules and materials rather than rigid inorganic materials commonly used in photolithography. Table 6.1 compares the features of photolithographic and soft lithographic approaches to surface patterning.83 In CP a PMDS stamp is used to transfer molecules of the `ink' to the surface of the substrate by contact, whereas in REM a PMDS polymer is used to duplicate structures of a master into multiple copies. In TM a drop of liquid pre-polymer is applied to the patterned surface of a PMDS mould. The filled mould is placed in contact with a substrate and irradiated or heated. After curing of the liquid precursor, the mould is peeled away to leave a patterned microstructure on the substrate surface. In MIMIC a PMDS mould in conformal contact with a substrate surface forms a network of empty channels

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Table 6.1 Comparison of photolithography and soft lithography approaches83 Photolithography

Soft lithography

Definition of patterns

Rigid photomask (patterned Cr)

Elastomeric stamp or mould (PDMS)

Materials that can be patterned directly

Photoresist

Photoresist

SAMs

SAMs Polymers Colloidal materials Sol-gel materials Organic and inorganic salt Biological macromolecules ca. 30 nm

The smallest feature size

ca. 50 nm

that can be filled with a pre-polymer. After curing the liquid into a solid, the PMDS mould is removed leaving a network of polymeric material on the surface of the substrate. SAMIM is a combination of replica moulding and embossing. A PMDS mould is wetted with a solvent, and is brought into contact with the surface of the polymer. The solvent dissolves a thin layer of the polymer, and the resulting gel-like fluid conforms to the surface topology of the mould. After evaporation of the solvent a relief structure is formed on the polymer surface. Soft lithographic techniques require little capital investment and are very simple. They can often be carried out under ambient laboratory conditions. They are not subject to the limitations determined by optical diffraction and optical transparency. Table 6.2 is a summary of soft lithographic and other nonphotolithographic approaches to creating patterned substrates.91,94,96-110 Producing patterned biomolecule layers using soft lithographic techniques has a specific importance to a number of growing techniques such as advanced tissue engineering, biomineralisation, DNA computing and cultured neural networks. It is promising for microfabrication of relatively simple, single-layer structures for use in cell culture, in sensors or as microanalytical systems.111,112

6.3.1 Microcontact printing (CP) (feature size obtainable down to 35 nm) CP is also a method for chemically and molecularly patterning surfaces down to sub-micrometre length scales. It has been used to pattern self-assembled monolayers (SAMs) of compounds such as hexadecanethiols and octadecytrichlorosilane (OTS) on gold113 and SiO2114,115 surfaces. CP has an advantage over conventional photolithographic patterning methods in that it requires no harsh chemicals making it suitable for patterning biologically active layers.

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Table 6.2 Summary of non-photolithographic methods83 Method

Resolution

Reference

Injection moulding Embossing (imprinting) Cast moulding Laser ablation Laser-induced deposition Electrochemical micromachining Ink-jet printing Stereolithography Soft lithography Microcontact printing ( CP) Replica moulding (REM) Microtransfer moulding ( TM) Micromoulding in capillaries (MIMIC) Solvent-assisted micromoulding (SAMIM)

10 nm 25 nm 50 nm 70 nm 1 m 1 m 50 m 50 m

96, 97 98, 99 100 101 102 103 104 105 106, 107 108 109 91 110 94

35 nm 30 nm 1 m 1 m 60 nm

Additionally, it should be possible to pattern multiple molecular layers by repeated application of CP. CP was mainly developed with SAMs of alkanothiol on gold113 and silver.116 The CP technique is very simple; a PMDS stamp is first inked with a solution of molecules, often proteins or alkanethiols, and brought into contact with the surface of a substrate. The soft PDMS stamp makes conformal contact with the surface, and molecules are transferred directly from the stamp to the surface in the time frame of a few seconds. After printing, unmodified regions can be back-filled by deposition of a second molecule. Figure 6.4 is a schematic representation of the CP process. A key requirement for successful CP is the appropriate surface chemistry. The surface chemistry should allow high spatial resolution, low background adsorption, and high selectivity for molecule immobilisation. These conditions can be accomplished by using SAMs. If suitable SAMs are chosen, marked contrast in properties such as wetting or protein binding may be obtained. For example, Prime et al. used CP to generate a pattern of hydrophobic alkanethiols sites where proteins deposited from solution while non-patterned regions were blocked with PEG to prevent protein adsorption.117 They extended this method by combining CP and chemical reaction of ligands containing both amino and anhydride groups providing patterned SAMs presenting several ligands on the same surface, especially polar, charged, or structurally complex groups such as proteins, polymers, or oligosaccharides.118,119 Patterned SAMs on gold can also be used as ultra-thin resists in selective wet etching, protecting coated areas from chemical etching, while uncoated gold areas are removed to leave a structured gold film.120 The ability to pattern SAMs by CP, and resulting control over the adsorption of adhesive proteins, enables the patterning of cells on substrates121 and enables change in the size and shape of cells.122

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6.4 Schematic of the well-established microcontact printing process developed by Prof. Georges Whitesides' group at Harvard University (taken from ref. 91, with permission).

However, this strategy cannot easily be extended to other non-metallic surfaces such as silicon or polymers. The importance of using patterning of silicon surfaces is immense for microelectronics, micro-machinery and micro-fluidics. Zhu et al. developed a two-step alkoxy monolayer assembly method for CP used directly on silicon surfaces based on a reaction between alcohol functional groups and Cl-terminated silicon surfaces.123 In an extended method they produced patterns of HO- and CH3O-terminated PEG regions whereby the HOterminated PEG monolayer was activated to immobilise protein molecules covalently.124 Generally, CP is inexpensive, very rapid and a powerful method for surface structuring, also applicable to curved substrates or inner surfaces and able to be used over a relatively large surface area (50 cm2) in a single step.125 Repeated printing using different stamps can allow complex surface patterns to be made of more than one kind of molecule. The smallest features generated to date with CP are trenches etched in gold with lateral dimensions of approximately 35 nm.126 By manipulation of the elastomeric stamps and the chemistry of SAMs formation, it is conceivably possible to reduce the size of features generated even further.127 Patterned microstructures fabricated using this technique may be useful in microelectronics, for micro-analytical systems, sensors, solar cells, and optical components. Patterned SAMs have also been used for the control of extracellular matrix protein adsorption and for the attachment of cells.128 It is possible to control the shape of the cell that attaches to a surface by creating islands of cell

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adhesion proteins onto substrate surface and thus to control cell growth and morphology. Polymers are widely used as biomaterials but cannot be modified by SAMs. Therefore, Chilkoti et al. developed a new method called microstamping on an activated polymer surface (MAPS) that enables patterning of biological ligands and proteins on polymer surfaces with a spatial resolution of at least 5 m.129,130 A requirement for MAPS is the presence of reactive functional groups in the polymer. Therefore, the surface of a polymer is modified to introduce the reactive group of interest. In a second step, an elastomeric stamp is inked with a biomolecule containing a complementary terminal reactive group and brought into contact with the activated polymer surface. This results in spatially resolved transfer and covalent coupling of the biomolecule on the polymer surface. To demonstrate the concept of MAPS, they patterned an amine terminated biotin molecule on carboxylated PET.131 Because MAPS is a simple and flexible method, it will be applicable to a wide variety of polymers that are amenable to surface modification. Hyun et al. demonstrated a simple method of patterning cells on commonly used polymeric biomaterials in the presence of cell adhesion proteins.132 This approach involves coating a polymer surface by solution casting a chemically reactive, biologically non-fouling comb polymer presenting short oligoethylene glycol side chains, activation of the surface, and patterning a cell adhesive peptide onto the surface of the comb polymer by MAPS to spatially direct the interactions of anchorage dependent cells with the substrate. The patterns of fibroblasts and endothelial cells were stable for long periods of time depending on the spacing distance between isolated features such as a square or stripe.133 Csucs et al. introduced a new variant of CP based on adsorption of a polycationic graft copolymer, poly-L-lysine-g-poly(ethylene glycol) (PLL-gPEG), on negatively charged surfaces through electrostatic interactions, rendering them highly protein and cell resistant.134 Furthermore, PLL-g-PEG can be modified by introducing biologically active ligands such as short cell adhesion peptide sequences (viz. RGD) at the terminus of PEG chains. Printing cell adhesive, peptide modified PLL-g-PEG followed with backfill with cell repulsive non-functionalised PLL-g-PEG, allows control of the concentration of biomolecules presented to cells at the interface. If the separation distance between adhesion sites was large (at least 30 m spacing) fibroblasts were found primarily on printed, cell-adhesive regions, whereas with smaller spacing between adhesive sites (less then 30 m) the cells were able to occupy nonadhesive regions as well. This observation can be explained by a bridging effect where cell pairs or agglomerates grow over non-adhesive regions without formation of cell-surface contact. Nanoscale objects such as colloidal particles can also be patterned on a variety of substrates including glass, silicon, and polymers using CP.135 An important feature of CP is that many proteins retain their biological activity after printing.136 It is possible to print the same surface with several

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proteins using different stamps having various inks printed many times or using parallel inking of a stamp followed by a single printing.137 However, two limitations of CP as presently used hinder its progress in printing arbitrary patterns of biological molecules. First, the PDMS elastomer used for the stamp in CP is hydrophobic. When printing with water-based, biological solutions, the poor wetting characteristic of the hydrophobic stamp yields extremely non-uniform patterns. Therefore, the hydrophilisation of PDMS is a requirement for employing biomolecules and polar inks for CP that do not have an affinity for native PDMS. The simplest method for preparing a hydrophilic stamp is oxidation of the PDMS surface by O 2 -plasma treatment,138,139 but the stamp does not remain hydrophilic for a long time because of the migration of low-molecular silicone residues from the bulk to the air-stamp interface.140 Delamarche et al. overcame this problem by grafting of a PEG silane onto plasma oxidised PDMS that resulted in a stable hydrophilic stamp surface for an extended period of time.141 The second limitation is that the soft PDMS stamp lacks the rigidity necessary for precision alignment and geometrical control of the pattern and due to stamp deformation is not suited for the printing of features below 250 nm. To resolve this problem, James et al. developed a thin elastomeric stamp on a rigid glass backing in order to allow printing of small isolated features and improve the alignment and geometrical control of the patterns.142 Another, often neglected, drawback of PDMS-based CP is the low molecular weight PDMS contamination presented on stamped surfaces.143 To overcome mechanical and contamination problems, Csus et al. investigated the possibility of new materials in CP applications. A copolymer of ethylene and a -olefin such as butene or octene (POPs) was used for printing protein patterns.144 Due to the higher bulk modulus of the polyolefin stamp, much higher printing quality in submicrometre ranges can be achieved with the POPs stamp than with the PMDS stamp. Also, multiple protein patterning is difficult to achieve; a method for printing different proteins onto a single substrate involves using stamps having different patterns or various inks that are printed many times onto the same substrate. It has been found that accurate reproduction of patterns realised in PDMS stamps on gold substrates is problematic on a scale smaller than 500 nm due to the diffusion of ink molecules from contacted to non-contacted areas.145 Li et al. used CP without ink, and obtained sub-micrometre edge resolution (less than 100 nm).146 They brought a surface oxidised PDMS stamp into contact with SAMs of acid labile adsorbates. The silicon oxide layer on the outer surface of the stamp was acidic enough to hydrolyse these adsorbates. Finally, scanning probe contact printing is a method for generating patterns less than 500 nm using a scanning probe with an integrated elastomeric tip to transfer chemical materials onto a substrate.147 Each contact printing action creates an individual dot. This method allows generation of structures made of organic monolayers on surfaces with excellent multi-ink and alignment registration capabilities.

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6.3.2 Micro-fluidic techniques (micro-moulding in capillaries ± MIMIC) In CP the deposition of liquid samples followed by drying can lead to poorly defined patterns, protein aggregation and the loss of biological activity. An alternative method to bring a controlled amount of biomolecule in contact with a surface without loss of biological activity is through use of a micro-fluidic network (-FN). The method is based on conformal contact ± a watertight seal that forms via van der Waals forces ± between a soft PDMS structure and a hard surface such as silicon, glass or some polymer. One or both of the surfaces is structured so that a network of fluidics channels is formed when the two surfaces are brought together. The channel network can guide solutions of proteins or other molecules over the substrate and the molecules adsorb onto the surface of the fluidics channels with the pattern defined by the micro-fluidic network (Fig. 6.5a).148 Once adsorption is complete, the PDMS is peeled off, to give a patterned surface with molecules deposited on surface regions exposed to the solution. Microfluidic networks are a very powerful tool for surface structuring. This method is well suited for patterning highly sensitive entities like proteins and cells on a variety of substrates, and can be used at room temperature with aqueous solutions. It is a simple technique that forms patterned structures in a single step with very high accuracy on a wide variety of surfaces ± flat or curved ± with structures as small as 100 nm. Different molecules can be made to flow in different channels of the network. Delamarche et al.110 used MIMIC to pattern a variety of substrates with immunoglobulins with submicron resolution (Fig. 6.6(a)Ð(c)). Only microlitres of reagent were required to cover several mm2 sized areas. This technique enabled simultaneous and highly localised immunoassay for the detection of different immunoglobulins. The independent network of capillaries allows simultaneous attachment of different biomolecules in each zone of flow, as shown in Fig. 6.6(b).149 Gradients can also be made using solutions at low concentrations. The high surface-to-volume ratio of the small channels results in a depletion of the solution as it travels through the network, resulting in lower and lower quantitites of molecules deposited on the channels walls. Patel et al. developed a micro-fluidic patterning technique for any biotinylated ligand to be patterned onto the surface of a biodegradable polymer. Spatial control over cell appearance was also observed when using these templates to culture endothelial and nerve cells. Furthermore, pattern features containing laminin peptide sequences were achieved for directional control of nerve regeneration.150 However, MIMIC cannot form isolated structures or patterns on contoured surfaces because it forms a hydraulically connected network of capillaries. Additionally, the extremely slow filling of small capillaries may limit the usefulness of MIMIC in certain types of nanofabrication

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6.5 The process of -FN for patterned delivery of proteins. (a) Patterned elastomer that forms a FN by contact with a substrate allows the local delivery of a solution of biomolecules to the substrate. (b) Flow of liquid between the filling pad and an opposite pad fills the array of microchannels that constitute the strategic part of this device. (c) Assembly of different zones of flow on the surface results from the independence of capillaries, each requiring only a small volume (1 l) of liquid to fill the zone and derivatise the underlying substrate. Left panel, top view; right panel, side cut along the channel (taken from ref. 110, with permission).

applications. Micro-fluidic networks must be sufficiently hydrophilic to promote filling of micro-channels by capillary action and should incorporate proteinrepellent surfaces to prevent depletion of proteins by adsorption to the walls of micro-channels. One approach to making micro-fluidic networks wettable and protein-repellent is to functionalise the channels with PEG.151 Another approach has been to prepare a polymer-based micro-fluidic device with defined and chemically reactive interfaces by deposition of sub-micrometre thin reactive coatings on the interior surface of micro-fluidic devices. The resulting coating was then used for immobilisation and self-assembly of a variety of biological ligands, proteins, and cells.152

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6.6 Schemes for the delivery and attachment of two different IgGs using a FN (a) followed by an immunoassay for the attached proteins after removal of the FN (b). A composite digital image shows light emitted from fluorescently tagged antispecies IgGs, each specifically recognising its binding partner previously patterned on a glass surface (c). The immunoassay is carried out with a heterogeneous solution of IgGs: tetramethyl rhodamine isothiocyanateconjugated antibody to chicken IgG (red), fluorescein isothiocyanateconjugated antibody to mouse IgG (green), and R-phycoerythrin-conjugated antibody to goat IgG (orange-red), each diluted 1:300 from their concentrated solutions. The left stripe comprises chicken IgGs and the right stripe comprises mouse IgGs. No light was evident from nonspecific deposition of antibodies to goat IgG anywhere on the surface. There was no green fluorescence on the left channel or red fluorescence on the right channel; each colour channel was collected independently so that such emission, resulting from cross-reactivity between the antibodies or their uncontrolled deposition, would have been easy to detect.

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6.4

Colloidal-based fabrication techniques

Colloidal particles as lithographic masks can increase the speed of patterning surfaces with nano-sized features. Colloidal particles of different materials can be produced with mono-disperse size distribution down to nanometre length scales. The colloidal particles served as an etch mask for the underlying substrate leaving the areas under the particles as hills. These particles can also be used as lift-off masks by filling up the surface surrounding the particles with a thin film. After removal of the particles, the underlying surface is exposed where the particles are located. One of the main advantages of this method over other lithographic techniques is that complex sample shapes can be patterned or coated with nano-porous layers. Spatz et al. have developed a copolymer micelle nanolithography method for generation of nanostructures on conductive and isolating substrates. The pattern dimension and geometry is controlled by the combination of the self-assembly of block copolymer micelles with pre-structures formed by photo- or e-beam lithography.153,154 This method allows bridging the length scale between several nm and 200 nm, through linking self-assembly nanostructures to structure sizes available from photo- or e-beam lithography. E-beam and photolithography by themselves are not able to write structures with such small dimensions over large areas, whereas self-assembly methods, such as the formation of nanoparticles in block polymer micelles cannot position particles in patterns with large ( m) separation distance and aperiodic arrangements. After formation of photo or ebeam resist patterning and coating with micelles loaded with a metal precursor, the substrate is dipped in a solvent to remove the resist film and any micelle deposited on top of the resist. The micelles inside the holes or grooves remain adsorbed at exactly the position where the hole originally formed. The substrate is treated with oxygen plasma in order to remove the polymer selectively, leaving behind Au-nanoparticles loaded into the core resembling the lithographic pattern. Combining these two techniques, gold dots of 1 nm diameter can be positioned with high accuracy. These dots can act as an anchor for individual macromolecules and biomolecules. The Au particles have been used to bind streptavidin proteins in an ordered array.155 Furthermore, a wafer with receptor modified gold dots may be used for the separation, location and screening of DNA or proteins. In order to group a small number of Au-nanoparticles, locally monomicellar layers were exposed directly to a focused electron beam so that a micelle located in an exposed area was chemically modified and immobilised on the substrate surface. This indicates that monomicellar layers can be applied as a negative ebeam resist. Au-nanoclusters are then deposited from diblock copolymer micelles by hydrogen plasma treatment. By this technique `micrometre nanostructured' patterns can be created marked by uniform 2, 5, 6, or 8 nm Au clusters.156

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Finally, colloidal lithography based on deposition of polystyrene particles onto flat oxidised titanium surfaces for particle sizes between 20 and 500 nm has been demonstrated. Methods to minimise aggregation with large particles or with high surface coverage has involved heating; co-deposition of silica particles and heating were shown to optimise the patterned protocols.157 Frey et al. developed a method termed ultra-flat nanosphere lithography (UNSL) that combines nanosphere lithography and ultra-flat template stripping to create 150 nm features of Au, Al, or Ag embedded in a matrix of Al, Au, or SiOx with topographical variation of less than 1 nm between patterned features and matrix.158 In UNSL, a material is deposited onto mica through a mask created by a close-packed monolayer of nanospheres. After removal of the spheres, a second material is deposited onto the nanostructures. Subsequently, upon removing the mica, the surface in contact with mica reveals flat nanostructures of the first material embedded in a matrix of the second deposited material. This technique enables the fabrication of chemically distinct patterns of one material embedded in a matrix of another material with minimal topographical variation. Independent control of chemistry and topography is important for studies of cell-surface interactions.

6.5

Template-imprinted nanostructured surfaces

The technique of molecular imprinting creates specific recognition sites in polymers by using template molecules.159,160 Molecular recognition is attributed to binding sites that complement molecules in size, shape and chemical functionality. Attempts to imprint proteins have been made with limited success.161,162 Shi et al. report a multi-step method for imprinting surfaces with protein-recognition sites using radio-frequency glow discharge plasma deposition to form polymeric thin films around proteins that were coated with disaccharide molecules. The disaccharides become covalently attached to the polymer film, creating polysaccharide-like cavities that exhibit highly selective recognition for a variety of template proteins.163 Figure 6.7 shows the fabrication process to create nanostructures or imprints that selectively adsorb individual proteins from a complex mixture.

6.6

Conclusion

As in physics and electronics, the multidisciplinary field of nanobiotechnology, which includes materials science, chemistry, biology and medicine, is becoming more reliant on the fabrication of structures and patterns with dimensions below 100 nm. To be useful in an industrial setting, ideally methods for nano-patterning and fabrication methods will need to be of low cost, flexible in terms of materials used, high precision during and after replication, be able to pattern on non-planar substrates, be produced with high

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6.7 Protocol for template imprinting of proteins. (a) Template protein was adsorbed onto a freshly cleaved mica in citrated phosphate-buffered saline (CPBS), pH 7.4. A 1±10 mM solution of disaccharide was spin-cast to form a 10±50 Ð sugar overlayer. The sample was put into the in-glow region of a 13.56 MHz RFGD reactor. Plasma deposition of C3F6 was conducted at 150 mtorr and 20 W for 3±6 min, forming a 10±30 nm fluoropolymer thin film. The resulting plasma film was fixed to a glass coverslip using epoxy resin and oven cured. Mica was peeled off and the sample was soaked in a NaOH/NaClO (0.5/1.0%) solution for 0.5±2 h for dissolution and extraction of protein. A nanopit with a shape complementary to the protein was created on the imprint surface. (b) A tapping mode AFM image of the surface of a fibrinogen imprint, together with a drawing of fibrinogen. (c) Mechanisms for the specific protein recognition of template-imprinted surfaces. A nanocavity-bound template protein is prevented from exchange with other protein molecules in the solution because of steric hindrance and an overall strong interaction; the latter is due to many cooperative weak interactions, involving hydrogen bonds, van der Waals forces and hydrophobic interactions for example (taken from ref. 163, with permission).

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speed and throughput and allow potentially for 3-D fabrication.164 Because all of these requirements may never be met by a single fabrication method, it may be necessary to create hybrid techniques which combine several of the methods described above. In particular, the present patterning techniques have proved to be well suited for specific applications, but do suffer from a number of limitations. Current photolithographic techniques that do not use a stamp require complex chemistry, while solvents used in conventional photolithography may denature or degrade the deposited layer. Lithographic techniques are costly because of the elaborate fabrication processes as well as the low manufacturing throughput. While nonlithographic techniques allow spatial control, the use of elastomeric stamps hinder reproducibility over large areas, transfer contaminants and degrade over time. Molecular self-assembly allows us to push the limits of soft lithography to the molecular scale. Similarly, creating boundary conditions by lithographic approaches allows direct self-assembly in more promising directions. The combination of both approaches, both powerful at their respective length scales, has large potential for future applications. In summary, there is great potential in the combination of self-assembly and conventional lithographic techniques. Self-assembly guarantees that the smallest sized features and functionalities are programmed into the structure and chemistry of molecules, while conventional lithography meets the demands of controlled structure location of nanostructures with large separation length and also aperiodic arrangements.

6.7

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Part II Measurement, monitoring and characterisation

7 Surface spectroscopies M M E H L M A N N and G G A U G L I T Z , University of Tuebingen, Gemany

7.1

Introduction

In recent years, some focus of research has been on biochemical and chemical sensors. The combination of physical sensors (transducers) with more or less analyte-selective layers of biochemical or chemical substrates has introduced selectivity to these systems. For this reason such set-ups have to be considered as complete sensor systems containing transduction principles, the sensitive layer, the signal processing, and evaluation strategies. Out of the huge variety of transduction principles, this chapter concentrates on optical techniques which provide many possibilities of application of optical principles. As this chapter is a review based on a lecture, a large number of optical principles will be classified and a survey on sensitive layers that differ as to sensitivity, selectivity, stability, and reversibility will be discussed. To cover the wide field in part, recent review articles have been included. This chapter is divided into three main sections. In the first section different surfaces for biomolecular sensing are discussed and examples of how to characterise these surfaces are given. In the second, different optical sensing methods are shown and compared with each other. In the last section some selected applications of biomolecular interaction analysis are given. The numerous publications in the area of optical sensing allow only the citation of some related publications and reviews.

7.2

Surfaces

A very large number of bioanalytical methods deal with monitoring interactions occurring at surfaces or at least uses surface bound ligands or receptors to monitor changes in analyte concentration even within a homogeneous phase. For all these detection methods there is a need for very special and smart surface properties. Selectivity, sensitivity, stability, and reversibility are examples of such requirements for these sensor systems which must be provided in part by the surface properties. In the case of biosensors, the user expects a rather high signal-

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to-noise ratio, short response times, low limits of detection, high sensitivity, and the possibility to use sensors also in real samples, not just in lab applications. Based on ion-selective electrodes, at the beginning, doped semi-conductor material was used as a sensing system. Later on layers derived from chromatography (Baldine and Bracci, 2000) were used as rather stable layers with high reversibility and very short response times. However these layers showed a quite low selectivity and especially for the detection of gases the limit of detection stayed only in the range of ppm. (Rathgeb and Gauglitz, 2000). Another approach to introduce selectivity to these layers is to use microporous material. According to the free volume of the micropores a discrimination by size is achieved (Lehner, 1996). Combining this principle with swelling properties of selected polymer films to detect gases or liquids allows even a discrimination by molecular dimensions and partition coefficients (Dieterle et al., 2002). For many years biomolecules have been used to provide very high selectivity and sensitivity. For example, antigen/antibody interactions, DNA/DNA hybridisation, inhibition of enzymes, protein/protein interaction and many other applications in the area of membrane bound receptors and signal cascades can be found. These biomolecules should exhibit a high receptor/ligand specificity, whereas polymers or organic sensitive layers exhibit non-specific interaction at higher reversibility. For all detection methods based on heterogeneous phase interactions nonspecific binding effects are a very important point. Therefore very sophisticated surface chemistry and modifications are a basic requirement to set up an efficient biosensor. One important approach is the silanisation of glass or quartz transducers with the subsequent covalent binding of various biopolymers supplying reduced non-specific binding properties and allowing functionalisation with ligands or receptors. This silanisation step can be characterised by NMR-spectroscopy and ellipsometry (Raitza et al., 2000). Dextran hydrogels (LoÈfas and Johnsson, 1990) supply a large number of functional sites within the volume (Piehler et al., 1996). However, in many cases, especially observing protein interactions non-specifity is not reduced enough. To solve this problems polyethylenglycole with different chain lengths bound to the silanised surface can be used as a shielding polymer against nonspecific binding (Feldmann et al., 1999). Ligands can easily be immobilised by means of either amino or carboxy functions. These layers are quite resistant against non-specific binding (Piehler et al., 2000). In contrast to the dextran hydrogel, the PEG-surfaces form a kind of two-dimensional polymer brush and interaction occurs only at the surface and not in the volume. Therefore these layers have a reduced number of interaction sites. Besides these principal approaches, a large number of other ideas have been realised, e.g., the use of avidin to immobilise biotinylated biochemical molecules to the transducer surface (Birkert et al., 2000), the use of His-tags

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(Gershon and Khilko, 1995) or the immobilisation of membrane structures of lipid double layers (Tien, 1985) to the transducer to model cell walls. All these various approaches have the intention to reduce non-specific binding, allow a large number of specific binding sites, increase the stability of the layer, which is essential for regeneration strategies, and increase selectivity and sensitivity. The disadvantage of such approaches is that due to the high binding constant the system is not reversible at all. Therefore regeneration strategies have to be introduced to regain the reusability of these sensing surfaces. One major drawback of these biomolecules is that although they increase the stability of the biolayer, their stability is not comparable with that of, e.g., polysiloxane films or microporous systems (Park et al., 1999). In order to try to combine the advantages of stability and reversibility with sensitivity and selectivity a wide field of research for many years has been supramolecular structures (Lehn and Ball, 2000) and biomimetic layers (Garnier, 2000). To increase selectivity of simple chemosensors supramolecular structures such as calixarene (Dickert and Schuster, 1995) were tried first. Another approach was the use of cyclodextrins (Schurig and Grosenick, 1994) or cycloheptapeptide (Jung et al., 1996) structures. Hybridisation studies were one of the first approaches to introduce selectivity by immobilising polynucleotide or peptide sequences. In the meanwhile, peptide nucleic acids (PNA, with peptides as a backbone) (Wang, 1999) have proven to be a better complementary binding system than DNA because of less repulsion by charges and better backbone stability, thus being stable against DNases and nucleases. Another approach is the synthesis of layers of molecular imprinted polymers (Haupt and Mosbach, 2000). During a co-polymerisation process a template is used to form a cavity with certain spacings and some selective binding sites. After removing the template the generated recognition and binding site is highly selective for the template (or molecules similar to the template) as a considered analyte. Although these layers are very stable and selective they often show very slow reaction times upon analyte exposure. To increase reaction times in the socalled spreader bar technique (Mirsky et al., 1999) the imprinting process is reduced only to the surface instead of the volume. Whether these layers will show potential in the future has yet to be proven. Figure 7.1 shows various different possibilities of assay formats. The best possibility would be a homogeneous assay which allows a direct interaction between the analyte and the reagent without any influence on surface properties. Especially in the case of optical detection, the problem is that this interaction has to change somehow the optical properties of the system. These might be simple colour changes. Therefore, in the case of homogeneous assays, mostly labelled systems are used. Although radioactive labelling allows a quite low limit of detection a fluorescence label is preferable. Therefore fluorescence labelling is normally used, taking advantage of either quenching effects or (in most cases) fluorescence resonance energy transfer.

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7.1 Assays using interaction between biomolecules in homogeneous phase and/or at heterogeneous interfaces. In both cases thermodynamics (equilibrium constant) and kinetics (association and dissociation rate constants) determine the interaction. Direct assays immobilise the receptor at the surface to measure the analyte, here a binding inhibition assay is demonstrated where derivatives of the analyte or ligand to detect are immobilised. In the preincubation phase receptor and ligand are mixed in the homogeneous phase, concentration of non-blocked receptor molecules is detected via the heterogeneous phase. High numbers of interaction sites at the transducer make this process diffusion controlled, at low `loading' the kinetics at the heterogeneous phase can be measured.

In the case of measurements at the heterogeneous phase two different detection principles can be distinguished, direct optical detection without any label and detection using fluorescence labels. In the case of direct optical detection preferably large analyte molecules have to be examined. The sensitivity of any kind of optical detection can be improved by increasing the analyte mass or volume. Therefore, either a competitive test scheme (labelled competing with non-labelled analyte) or a so-called binding inhibition test scheme is used. This represents the following assay type: an analyte derivative is immobilised in the biopolymer; in a preincubation phase the analyte and the reagent are mixed together; the analyte as a ligand blocks receptor sites, which cannot react any more in this blocked state with the analyte derivative immobilised to the surface. This approach can be realised with a labelled reagent or direct optical applications.

7.3

Optical detection methods

Monitoring of effects at surfaces or interfaces using regular reflectance of light can be divided into two basic detection principles based on reflectometry or refractometry. Both use the influence of the thickness and/or refraction index of the observed layer on the phase and/or amplitude of electromagnetic radiation

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7.2 Schematic drawing of regular reflection, resulting refraction and total internal reflection for angles larger than the critical angle in the waveguide. The guided wave exhibits an evanescent field close to the waveguide, which is influenced by absorbing molecules or changes in the refractive index within the penetration depth of this evanescent field. Accordingly, the coupling of this external field to the guided field vectors within the waveguide influences the effective refractive, the transversal electric (TE-) and magnetic (TM-) differently (Hecht and Zajec, 1980).

penetrating this layer or being reflected. Different detection principles are schematically demonstrated in Fig. 7.2. In the case of refractometry, radiation guided in a wave guide is influenced by minimum changes in the refractive index or transmittance of an adjacent medium, since its evanescent field probes this medium resulting in an effective refractive index. Thus the TE (transversal electrical) and TM (transversal magnetic) modes of the wave propagating in the wave guide are influenced differently. A review of a variety of different optical detection principles based on evanescent field techniques can be found in Gauglitz (1996). Figure 7.2 demonstrates that incident radiation is reflected in part at an interface between media with different refractive indices according to the Fresnel equations. Radiation passing through the medium with high refractive index is refracted to the optical axis (Fig. 7.2, left side). Radiation passing through the interface in the opposite direction will be refracted away from the optical axis. Thus, a critical angle exists beyond which radiation will not pass through the interface out of the medium of high refractive index; it will be `totally reflected'. Accordingly, under such conditions, radiation will be guided in this medium being a fibre or, more generally, a waveguide (planar or stripe). The electric field of this guided wave causes an evanescent field existing close to the interface in the medium of lower refractive index. The evanescent field penetrates the adjacent medium just by approx. half of the wavelength of the

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guided radiation and also shows an exponential decay. Therefore, any changes in these properties influence the evanescent field and accordingly the propagation of radiation in the waveguides. The `effective refractive index' of the waveguide varies. Interferometer type waveguides like Mach-Zehnder interferometer (Ingenhoff et al., 1993) or Young interferometer (Brandenburg et al., 1992) use two different waves guided in two different waveguides. One of these waveguides is exposed to the analyte while the other is used as a reference channel. This leads to a detectable difference in phase of the two waves. For the Young interferometer the waveguide arms are not reunifed but rather image the interference pattern produced by the two open ends of the waveguide arms on a CCD (Brandenburg et al., 1992). Using both TE and TM modes allows internal referencing. This has been improved by Lukosz and Stamm (1991) in its mode beat interferometer, measuring amplitudes and phases of both polarisation states. Another device that is frequently used for monitoring changes in refractive index at surfaces is the grating coupler (Clerc and Lukosz, 1994). The grating coupler combines a waveguide layer with layers in which a grating is embedded. The grating constant is influenced by the refractive index within the adjacent medium. Similar to an interference filter, this grating condition varies with the angle of incidence or determines the preferred wavelengths. Dependent on the effective refractive index (that is strongly influenced by the adjacent layer where the interaction takes place), wavelength and angle of incidence radiation incident to the grating will be reflected or coupled in the waveguide (Nellen and Lukosz, 1993). The reflected radiation is monitored using either an angle resolved set-up or a CCD camera (avoiding mechanical parts). A special set-up uses so-called bi-diffractive couplers where two different grating constants are superimposed resulting in a different angle for the outcoupled wave compared to the direct reflected one. Another type of interrogation of the polarisation status is applied in prism couplers (Cush et al., 1993). The radiation couples out of a prism via the frustrated total internal reflection of a low refractive index layer (with a thickness of ~1000 nm) into the high refractive index waveguide. A 45ë polarisation is chosen, and TM and TE modes travel in the resonant layer (waveguide: thickness of ~100 nm), differently influenced via evanescent field by changes in the adjacent medium. Thus, the polarisation state changes in this `resonant mirror'. The angle at which this coupling occurs, the resonant angle, is, essentially, dependent upon the refractive index at the surface of the sensor. Changes in refractive index of the adjacent layer will change the resonant angle. The best researched evanescent field technique is surface plasmon resonance being reviewed in theory and application to chemo- and biosensing in many articles (Homola et al., 1999, 2002). A prism is coated at its basis by an approx. 50 nm metal film. When the energy of the incident photon electrical field is just right it can interact with the free electron constellations in the metal film. The

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incident light photons are absorbed and converted into surface plasmons resulting in a `dip' in the reflectance diagram. The resonance condition of these plasmons depends, via the evanescent wave, also on the refractive index at the surface opposite to the waveguide. Simplified, the binding of biomolecules results in the change of the refractive index on the film, which is measured as a change in reflected light. This direct optical detection method has been commercialised the longest. In recent years, additional reviews have been published regarding SPR techniques (Rich and Myszka, 2000; Van Der Merwe and Anton, 2001; Davis and Wilson, 2001; Sadana, 2001), resonant mirror (Kinning and Edwards, 2002), grating coupler (Voros et al., 2002; Kuhlmeier et al., 2003), Bragg gratings (Santos and Ferreira, 2002).

7.3.1 Reflectometry At the interface of a thin layer the incident light is partially reflected (depending on the difference of refractive indices of the two layers) whereas the other part penetrates the layer and is again partially reflected at the second interface. These two partially reflected beams can superimpose and form a characteristic interference pattern depending on the wavelength, the angle of incidence and the refractive index and the physical thickness (the so-called `optical thickness') of this layer. Changes on or at this layer (for example, due to analyte binding to this layer) result in a characteristic shift of this interference pattern that is easily detected. The principle of this simplified version of ellipsometry as shown in Fig.7.3 is called reflectometric interference spectroscopy (RIfS). It is a simple and robust technique in chemo- and biosensing as demonstrated later in the application section. In ellipsometry the polarised light is used to probe the dielectric properties of a sample. The most common application of ellipsometry is the analysis of very thin films. Through the analysis of the state of polarisation of the light that is reflected from the sample, ellipsometry can yield information about layers that are thinner than the wavelength of the light itself, down to a single atomic layer or less. Depending on what is already known about the sample, the technique can probe a range of properties including the layer thickness, refractive index, morphology, or chemical composition. The first introduction of ellipsometry was as far back as the 1940s (Azzam and Bahara, 1988; Arwin and Aspnes, 1986), and today it has many standard applications. It is mainly used in semiconductor research and fabrication to determine properties of layer stacks of thin films and the interfaces between the layers. However, ellipsometry is also becoming more interesting to researchers in other disciplines such as biosensors and medicine. These new disciplines require new techniques such as microscopic imaging. The polarisation state of the light beam can be determined in many different ways. In the nulling technique the polarising elements were rotated until the

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Surfaces and interfaces for biomaterials

7.3 The principle of reflectometric interference spectroscopy (RIfS) is based on white light interference according to the given formula. Changes in the optical thickness of the layer between the two interfaces causes a change from second reflected beam I2 to I0 2 superimposed to I1 and from a destructive interference (at a specific wavelength: broken line) to a positive constructive interference. Considering the whole interference pattern a shift is observed correlating to the amount of change in physical thickness. IR can be used as a reference. The change in physical thickness is caused either by swelling of a polymer layer (uptake of analyte) or by an affinity reaction adding biomolecules to receptor molecules at the interface.

signal at the detector is extinguished (nulled). This technique was also used in the first ellipsometers. The measuring of a signal near `zero' is one disadvantage of this nulling technique, as modern light detectors have a significantly higher noise at low signal intensities. Therefore other techniques to determine the polarisation state, e.g., phase modulated ellipsometry, have been developed. Both principles (reflectometry and refractometry) measure changes in the optical thickness of a sample. However, it has to be considered that the refractive index is a highly temperature dependent value. Thus all devices based on evanescent field techniques need a good thermostating (below 0.1 K) or need to be referenced well using a dual-channel instrument. However, in the case of reflectometry thermostating normally is not required. By chance an increase of the temperature leads to an increase of the thickness of the layer due to thermal volume expansion. This increase results in a decrease of the refractive index of the layer (the increase of the layer leads to a lower density of polarisable molecules and thus to a lower refractive index). These two effects nearly compensate each other.

7.4

Biomolecular interaction analysis

In principle, monitoring of interactions with label-free biosensors can be divided into two assay formats. First, the interaction between the immobilised ligand and

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7.4 Scheme of a simple binding experiment. Under appropriate experimental conditions the association and the dissociation rate constants, ka and kd, and also the affinity constant K can be determined.

the dissolved binding partner can be monitored directly, enabling the study of kinetics and the thermodynamics of the (heterogeneous) reaction. This direct detection is used frequently for macromolecules like proteins. An indirect assay format is often applied for the detection of interactions between a macromolecule and a low molecular weight species. In this type of assay both examined binding partners are in solution and the equilibrium concentration is quantified with the biosensor, allowing the calculation of the thermodynamic properties of the `homogeneous' reaction (Fig. 7.4).

7.4.1 Direct detection The easiest case is a simple 1:1 stoichiometric reaction according to: A‡Bˆ

ka kd

7:1

where A is the immobilised binding partner onto the transducer surface, B the soluble binding partner, and ka and kd are the association and the dissociation rate constants, respectively. Assuming a constant concentration of the soluble

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binding partner (for example, in a constant flow device) the binding reaction can be described as a pseudo-first-order reaction according to the Langmuirian theory: dÿ=dt ˆ ka C0 …ÿmax ÿ ÿ† ÿ kd ÿ

7:2

where ÿ is the surface coverage of captured ligate, C0 is the solution concentration of ligate, and ÿmax is the total binding capacity of the surface. Assumptions made for the validity of this model are: the binding event is controlled only by the intrinsic kinetic rate constants ka and kd and the diffusion of the soluble binding partner to or from the surface is not rate-limiting, the concentration of the ligate C0 is constant; the interaction of the biomolecules occurs in 1:1 stoichiometric ratio. Either by applying a linear transformation to eqn 7.2 according to Karlsson et al. (1991) or by an exponential curve fitting as proposed by O'Shannessy et al. (1993) the corresponding values for ka and kd can be obtained. The dissociation rate constant kd can also be calculated from the dissociation of the bound ligate from the surface during the dissociation (wash out) phase by replacing the ligate solution by buffer. The dissociation can be described by the first-order equation: dÿ=dt ˆ ÿkd ÿ

7:3

Even if the methods of linear transformation are simple and practical to use they imply the disadvantage of error propagation by the transformation and suppress the fact that the data are often non-linear. Therefore it is clear that this kind of analysis gave rise to criticism (O'Shannessy, 1994). The benefit of an exponential data analysis is that the errors of the kinetic rate constants directly reflect errors in the primary data and the curve fitting is permanent under the control of the operator. Moreover the values for ÿeq (the extrapolated equilibrium response) and ÿmax are fit which is possible even if the binding curve does not reach the equilibrium within the measurement time. The third method to derive kinetic rate constants is also an iterative curve fitting routine called global analysis (Roden and Myszka, 1996). This method is applied simultaneously to complete sets of binding curves at varied conditions. Using a reference channel the data can be corrected for drift, non-specific binding and refractive index effects. A comparison of the three analysis routines, linearization, integrated rate equation and numerical integration is given in Morton et al. (1995). A major problem of all these analysis methods is the possible difference between the assumption of a first-order reaction limited only by reaction kinetics to the real experimental data. Significant deviations from ideal conditions will lead to wrong results. For example Nieba et al. (1996) found large differences between dissociation constants derived from kinetic rate constants measured at the surface and determined in the solution.

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Possible differences between experiment and model may be a non-negligible influence of mass transport to the surface, an underestimation of the dissociation constant due to rebinding effects or a more complex binding mechanism that could not be described sufficient by eqn 7.1. A detailed description of the influence of mass transport on the data can be found in Glaser (1993). The most important parameters are the product ka0 g and the Onsager coefficent of mass transport Lm, with ka0 as the observed associate rate constant and g the surface concentration of free binding sites. The Onsager coefficient is according to SjoÈlander and Urbaniczky (1991) dependent on the geometrical dimensions of the flowcell r 2 3 D f Lm ˆ km ˆ 0:98
7:4 h2 bl with the diffusion coefficient D, the flowrate f, the channel height h and its width b and the distance l form flow cell entrance. According to Glaser three cases can be distinguished: 1. 2. 3.

ka0 g  Lm , with no mass transport influence on the kinetics. ka0 g and Lm are in the same order of magnitude. Then the true rate constants ka and kd are underestimated up to several orders of magnitude. ka0 g  Lm with full mass transport control. No kinetic information can be gained.

Methods for data evaluation of mass transport limited binding kinetics are suggested in several publications (Schuck, 1996; Myszka et al., 1996; Schuck and Minton, 1996). For data analysis, very often the association phase is split into the mass transport limited phase and the kinetically limited phase. This can be carried out by the determination of the linear part of a dCt/dt versus Ct plot. However, especially for high association constants this method underestimates the values of ka. Schuck and Minton (1996) suggested a two-compartment model taking mass transport into account and considering the transport only from the bulk compartment to the surface. However, the best way to resolve this problem is to reduce the mass transport effects by special experimental conditions. According to eqn 7.4 the Onsager coefficient can be influenced by using appropriate flowrates and cell geometries, and the reduction of the surface concentration of free binding sites g will reduce the influence of mass transport (the apparent rate constant ka0 is a fixed constant for a given biochemical system and cannot be changed). On the other hand the reduction of the binding capacity results in lower signals. Thus a high sensitivity and low S/N ratios of the biosensor equipment are required. These parameters determine the upper limit of the range of association rate constants amenable to study by optical biosensors (Hall et al., 1996). Another reason for differences between experiment and the simple model according to eqn 7.1 is the so-called rebinding. This rebinding results in an

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Surfaces and interfaces for biomaterials

underestimation of the dissociation rate constant kd. The probability for the rebinding to the surface increases with the concentration of free binding sites (Shoup and Szabo, 1982). This influence can also be reduced by lowering the surface concentration of free binding sites g, or by adding a competitive binder in high concentration to the washing buffer, to prevent dissociated molecules rebinding to the surface. The (heterogeneous) thermodynamic constant KA can be calculated either from the kinetic rate constants ka and kd or by evaluating the concentration dependent equilibrium binding signals. For 1:1 reactions the Langmuir adsorption isotherm is: ÿeq ˆ ÿmax

C0 1 ‡ C0 KA

7:5

A Scatchard Plot ÿeq =C0 versus ÿeq results in a straight line with the slope ÿKA.

7.4.2 Indirect detection The constants measured at the surface often differ from those in the solution. Therefore a different approach to determine the homogeneous association constant KA has been developed using the sensor surface to probe for the concentration of unbound analyte molecules in solution. The principle is given in Fig. 7.5. An equilibrated mixture of analyte and the corresponding binding partner (ligate) is injected over the sensor surface where an analyte-derivative is immobilised on it. From the initial binding rate the concentration of unbound ligate in solution can be determined (Karlsson et al., 1994; Edwards and Leatherbarrow, 1997). According to eqn 7.2, for the beginning of the binding (ÿ ˆ 0 pg/mm2): dÿ=dt ˆ ka C ÿmax

7:6

the signal is directly proportional to the concentration of unbound ligate C. If the reaction is under a pure mass transport control the initial binding rate can be described according to Fick's first law: dÿ 7:7 ˆ km
C dt with km as the diffusion coefficient. Therefore in both cases the initial binding rate is directly proportional to the concentration of unbound ligate in solution. A binding rate proportional to concentration of free ligate C is always given if the amount of bound ligate does not exceed 10% of the maximum binding capacity, therefore high binding capacities of the surface are desirable (Piehler et al., 1997a).

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7.5 Probing of the equilibrium concentration of unbound ligand with the surface-bound ligand.

For a known initial concentration of ligate C0 the dissociation constant can be derived from: c
a …C0 ÿ x†
…A0 ÿ x† ˆ 7:8 x x with c and a as the equilibrium concentrations of ligate and analyte, and C0 and A0 as the corresponding initial concentrations, and x as the equilibrium concentration of formed complex. Equation 7.8 can be rearranged to: s C0 ÿ A0 KD …C0 ‡ A0 ‡ KD †2 ÿ C 0 A0 ‡ cˆ 7:9 4 2 KD ˆ

The resulting secondary plot (titration curve) for different concentrations of analyte C0 is shown in Fig. 7.5. By fitting eqn 7.9 to the titration curve the dissociation constant KD can be determined. This kind of assay is often denoted as a competitive format, but this term implies that a competition between ligate,

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Surfaces and interfaces for biomaterials

(immobilised) analyte derivative and analyte (dissolved) occurs. Due to the short time of interaction in a flow system between the surface and the solution of ligate and analyte no equilibrium for these three species can be reached (in fact the equilibrium in the solution is not affected by the surface if dissociation of the ligate-analyte complex is not too high) and therefore the binding rate depends only on the concentration of free ligate. For this reason the term inhibition test is a better choice. The advantages of this assay type are that due to the indirect assay format also small analytes can be investigated, and that the effects caused by the immobilisation (rebinding, mass transport) do not affect the results. Also many different analytes can be examined using the same surface. Recently, even a new detection principle has been developed. A time resolved observation of the interaction process during the incubation step allows even the determination of the on and off-rates (ka and kd) for the homogeneous reaction (between ligand and dissolved analyte). In this case, the surface is used only to determine the concentration of unbound ligand in solution. Due to steric hindrance, for example, these on and off-rates can be quite different compared to the on and off-rates for the heterogeneous reaction (between ligand and surface bound analyte).

7.5

Conclusion

Optical sensors have proven in the past to be either very simple and cost-effective devices or to allow rather sophisticated multisensor applications. As a large number of different optical principles exist, in principle many of these methods can be applied to a huge number of applications. It has to be noted that in most cases the choice of an appropriate assay design, and therefore the surfaces and transducers used are linked together very closely and cannot be considered separately. It becomes evident that none of the afore-mentioned various sensing principles, and also the unmentioned electrochemical or mass sensitive devices, is superior in general. Their feasibility is strongly dependent upon the application they are used for. This became obvious when comparing various refractometric and reflectometric methods in the same biomolecular interaction study using antibodies and antigens, and setting up the surface chemistry by the same person. In the case of both studies, the limits of detection ranged for all methods examined within one order of magnitude. The discrimination was achieved only at the cost of expenditure on apparatus and by the sophistication of the fluidics used (HaÈnel and Gauglitz, 2002; Piehler et al., 1997b). The essential result of these considerations is certainly that research in the area of sensing requires interdisciplinary understanding of the detection principles, of the sensitive layer, of the kinetics and thermodynamics of interaction processes and of the fluidics. Thus fundamental research has to characterise these layers and the interaction processes to improve the understanding which is the prerequisite of any optimisation approach.

Surface spectroscopies

7.6

197

References

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Park J, Groves W A, Zellers E T (1999), `Vapor recognition with small arrays of polymer-coated microsensors. A comprehensive analysis', Anal Chem, 71, 3877. Piehler J, Brecht A, Geckeler K E, Gauglitz G (1996), `Surface modification for direct immunoprobes', Biosens Bioelectron, 11, 579. Piehler J, Brecht A, Giersch T, Hock B, Gauglitz G (1997a), `Assessment of affinity constants by rapid solid phase detection of equilibrium binding in a flow system', J. of Immunol. Meth., 201(2), 189±206. Piehler J, Brandenburg A, Brecht A, Wagner E, Gauglitz G (1997b), `Characterization of grating couplers for affinity-based pesticide sensing', Appl. Opt., 36, 6554. Piehler J, Brecht A, Valiokas R, Lidberg B, Gauglitz G (2000), `A high-density poly(ethylene glycol) polymer brush for immobilization on glass-type surfaces', Biosens Bioelectron, 15, 473. Raitza M, Herold M, Ellwanger A, Gauglitz G, Albert K (2000), `Solid-state NMR and ellipsometric investigations of C30 chains bonded to SiO2 surfaces', Macromol Chem Phys, 201, 825. Rathgeb F, Gauglitz G (2000), `Optical gas sensors in analytical chemistry: applications and trends and general comments', in Encyclopedia of Analytical Chemistry, ed. R A Meyers, John Wiley & Sons Ltd, Chichester, 2189. Rich R L, Myszka D G (2000), `Survey of the 1999 surface plasmon resonance biosensor literature', J Mol Recogn, 13, 388. Roden L D, Myszka D G (1996), `Global analysis of a macromolecular interaction measured on BIAcore', Biochem Biophys Res Com, 225, 1073-1077. Sadana A (2001), `Kinetic analysis for analyte-receptor binding and dissociation in biosensor applications: a fractal analysis', Biotech Genetic Eng Rev, 18, 29. Santos J L, Ferreira L A (2002), `Fibre Bragg grating interrogation techniques', in Handbook of Optical Fibre Sensing Technology, John Wiley & Sons, Chichester, 379. Schuck P (1996), `Kinetics of ligand binding to receptor immobilized in a polymer matrix, as detected with an evanescent wave biosensor. I. A computer simulation of the influence of mass transport', Biophysical J, 70, 1230±1249. Schuck P, Minton A P (1996), `Analysis of mass transport-limited binding kinetics in evanescent wave biosensors', Anal Biochem 240: 262±272. Schurig V, Grosenick H (1994), `Preparative enantiomer separation of enflurane and isoflurane by inclusion gas chromatography', J Chromatography, A666 617. Shoup D, Szabo A (1982), `Role of diffusion in ligand binding to macromolecules and cell-bound receptors', Biophys J, 40, 33±39. SjoÈlander S, Urbaniczky C (1991), `Integrated fluid handling system for biomolecular interaction analysis', Anal Chem, 63, 2338±2345. Tien H T (1985), `Planar bilayer lipid membranes', Prog Surf Sc, 19, 169. Van Der Merwe, Anton P (2001), `Surface plasmon resonance', in Protein-Ligand Interactions: Hydrodynamics and Calorimetry, Oxford University Press, 137. Voros J, Ramsden J J, Scucs G, Szendro I, De Paul S M, Textor M, Spencer N D (2002), `Optical grating coupler biosensors', Biomaterials, 23, 17, 3699. Wang J (1999), `PNA biosensors for nucleic acid detection', Curr Issue Molec Biol, 1(2), 117.

8 Surface microscopies C Z I E G L E R , University of Kaiserslautern, and Institut fuÈr OberflaÈchen- und Schichtanalytik GmbH, Kaiserslautern, Germany

8.1

Introduction

Biomaterials must meet the demands of materials science on various length scales as well as clinical requirements. In particular, the interface is important for biomaterials, as it defines the interaction with the environment,1,2 as described elsewhere in this book. Interface structures are in dimensions of a few nanometers for proteins up to the micrometer range for cells in the field of biomaterials. In Fig. 8.1 an overview of the interfaces on different scales is presented to show the parameters that can be of interest. Macroscopically, the shape of an implant, the structure, and its mechanical stability are important. The microscopic level is determined respectively by morphology (i.e. domain structure, the presence of ionic groups and chemical composition respectively, including further modification), topography (i.e. surface roughness, planarity, feature dimensions), hardness and elasticity (Young's modulus). These characteristics determine other features like wetting behavior and interaction forces (inter- and intramolecular) including cell-surface or cell-cell interactions. Various degrees of information about these properties can be obtained using different analysis methods including microscopic and spectroscopic methods. Mechanical, biochemical, and optical properties depend mainly on the topography and chemistry of the surface. Topography may be defined as size, shape, distribution, and the hierarchy of surface features. These features are either discrete (holes and peaks) or continuous (furrows and ridges) in random (statistical), fractal (self-similar on different length scales) or periodic distribution across the surface. Rough surfaces exhibit a larger surface area, and more contact points for biological molecules and cells. The reason for larger entities like cells to adhere to the surface and how they arrange their layer growth is dependent on the size and distribution of the surface features. The chemical properties of the interface between biomaterial and body environment determine the interaction of the surface with water molecules, ions, biological macromolecules and cells. Surface reactivity depends on chemical composition,

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8.1 Different relevant length scales in biomaterials research demonstrated for the example implant. The biomaterial-tissue interaction is based on molecular events, which affect meso- and macroscopic material properties. On larger scales additional (e.g. cellular, mechanical) effects resulting from combining individual units (molecules, cells, crystallites) to a more extended ensemble arise.1

the production process and pretreatment before use of the materials, and is essential for fixation, growth and proliferation of tissue and bone cells. Cell adherence via membrane receptors and adhesion proteins is influenced by active surface groups (generated, e.g., by oxidation of metals or coating of materials with specially designed polymer layers) which regulate the adsorption of the anchoring protein layer. The functionalization of the substrate additionally controls wetting behavior where particularly hydrophilic surfaces show enhanced cell adhesion. The latter is due to the fact that systems like cells

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and hydrophilic surfaces reach a thermodynamically favorable state by combining their high energetic surfaces. In addition, even hydrophobic interfaces like gold can get more hydrophilic by protein adsorption,3 which may help to increase biocompatibility. Polarizability and charging of the interface affect electrostatic interactions between charged species and the surface. Charge screening and complexation by multivalent ions present in all types of body fluids allow attraction between molecular and surface groups of like charge, and stabilize biological layers. At conductive interfaces, electrochemical reactions with charge transfer between the electrolyte solution and the substrate may occur, which interfere with cellular metabolism and the conformation of adsorbed adhesion proteins. This can set free toxic substances and cause allergic and inflammatory body reactions. Topographical as well as chemical effects play a role in the tribological properties of biomaterials. Tribology describes the behavior of interfaces in motion, and becomes important when implants are designed to support body movement. In this case, friction, wear, and lubrication of implant and body respectively, at different moving implant elements, influence the function and longevity of the applied biomaterial. Interfacial friction depends on the type and strength of interaction forces, the clasping of surface elevations and troughs, and the force transfer properties of intermediate fluids. Hardness, cohesion, and adhesion of individual surfaces in contact determine to what extent materials are worn down by abrasion (e.g. scratch and particle formation), adhesion (e.g. welding processes) and surface fatigue (e.g. crack formation). Liquid films in the crevices between two hard surfaces can reduce friction and wear. The effect of such lubricants is influenced by surface chemistry and separation, as well as by the viscoelastic, and hence force transferring, properties of the fluids themselves. This short discussion of biomaterials in relation to surface properties shows that the interplay between implant and body environment can be very complex, and includes a large variety of parameters from materials science, biology, and medicine, explained in more detail elsewhere.4 The need for surface analysis, in particular with nanometer resolution, arises on the one hand from the importance of the interface between natural and artificial materials, and on the other hand from the high dependence of macroscopic behavior on the microscopic features of the biomaterial. The following sections will show the different capabilities of today's microscopic techniques, and the criteria by which these methods could be used for certain problem-solving strategies.

8.1.1 Different concepts of imaging surfaces Imaging techniques can be categorized under different concepts. In most of the methods, probes such as beams of electrons, ions, or photons will come into

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8.2 Schematic presentation of typical arrangements to study solid surfaces.

contact with the sample surface. After interaction, either the beam of these primary probes will be analyzed for its changed properties, or secondary probes which are released from the sample surface are detected (Fig. 8.2). There are two principal types of techniques to obtain a real image of the sample. For the first one (Fig. 8.3) the probe beam has to have a small diameter and the sample surface can then be scanned point by point. The software then constructs an image from these pixels. The classical type of such an imaging instrument is the scanning electron microscope (SEM). A particular form of scanning techniques are the so-called scanning probe methods such as scanning tunneling microscopy (STM), scanning force microscopy (SFM), and scanning nearfield

8.3 Scanning techniques as one possible measuring principle to characterize geometric structures.

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optical microscopy (SNOM). Alternatively, the surface is 'illuminated` by the whole beam, and through the optical path of the instrument, a direct image can be obtained. Important techniques are the optical microscope and the transmission electron microscope (TEM). Diffraction techniques which give indirect information of the geometric structure of highly ordered materials by determining the reciprocal lattice structure will not be discussed here because biomaterials are usually not single crystalline materials.

8.1.2 Imaging parameters and requirements of the different methods To choose the right microscopic technique, one has to evaluate what information is required from the sample surface. Important questions to be answered are: · What is the required lateral resolution to see the effect? · Do I only need a topographic image of the surface, or do I need additional information, such as elemental composition, mechanical properties, or optical properties with microscopic resolution? · How surface sensitive do I need to be? Is this property present only at the surface or throughout the bulk? Do I need information on the depth to which this property occurs, or is it important only to see whether it is present or not? · Are there requirements for the sample environment, i.e., do I expect changes in properties if the sample is put into vacuum or is covered with metal etc.? Lateral resolution The lateral resolution of scanning techniques is determined by the larger diameter of the following two effects: probe beam diameter and surface area in which the analyzed information is generated. To show the difference, the interaction of electrons with a sample surface is shown in Fig. 8.4.5 The probe beam of primary electrons has a typical diameter of a few nanometers. By impacting on the surface, the primary electrons can be reflected (reflected or backscattered electrons, RE), create secondary electrons (SE) of different origin or produce X-rays. The two latter points will be explained in more detail in section 8.3. From Fig. 8.2 one can see that the SEd, which are directly produced by the primary electrons, come from an area of about the beam radius, i.e., typically 1±10 nm. The RE, the SEid, which are indirectly produced by the RE, and a special type of SE, the Auger electrons (see section 8.3) come from an area of about 1 m. Therefore the lateral resolution by imaging REs and SEid (for practical solutions see section 8.3) is in the micrometer range, whereas by imaging SEd a resolution of 10 nm or better (down to 1 nm) is possible.

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8.4 Backscattered (RE) and secondary electrons (SE) in an electron microscope. For details see text. KL are the characteristic X-rays, zSE the information depth of the secondary electrons, zA that of the Auger electrons, zRE of the backscattered electrons.5

In the far field, i.e., if the detector is far away from the sample if compared to the wavelength, the resolution of direct imaging techniques is limited mainly by the wavelength of the utilized probes because of diffraction. According to Abbe, the resolution of an imaging system depends on the wavelength , the diffraction index n of the volume between sample and objective, and the angle  under which the light passes through the objective:  8:1 n
sin  The term n
sin  is called the numerical aperture. It can be varied only to a small extent, so the resolution of light microscopy with typical wavelengths around 500 nm is limited to 200 nm. Electrons can have wavelengths in the Angstrom region, therefore atomic resolution is possible in principle. In electron microscopy, the glass lenses have to be replaced by electrical and magnetic lenses, which deflect the electron beam in an analagous way to glass lenses with A ˆ 0:61

x

x (x)

TEM

STM SFM Optical methods

SEM

Atomic resolution

x x

Height profiles x (by Auger or by EDX) x (by EELS or by EDX)

Elemental composition

Table 8.1 Measurable parameters of important imaging techniques

x (by labeling)

Specific biological groups

x

Hardness

x

Elasticity

x

Friction

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light. Because of the aberration of these electromagnetic lenses, the resolution in electron microscopy is limited to 0.2 nm. In the so-called near field, the classical diffraction limit is no longer valid (see section 8.5). Atomic resolution is therefore theoretically possible by all scanning probe techniques. However, only STM gives easily atomically resolved images; SFM shows atomically resolved structures usually only under ultrahigh-vacuum (UHV) conditions and at low temperatures, and so far no atomically resolved SNOM images have been observed due to technical problems (compare section 8.5). Additional information As outlined in section 8.2.1, there are many different properties which are of interest to fully characterize a biomaterial. Therefore, it is important to know which information can be obtained by which method. This is summarized in Table 8.1, and will be explained briefly in the respective sections below. Surface sensitivity Information depth is determined by the so-called mean free path of the primary and/or secondary probes. The mean free path is defined as the mean distance the probe can travel before it undergoes a scattering event. Scattering can result in energy and/or momentum change of the probe. Because interaction of charged particles such as electrons and ions is much larger with the atoms of the sample surface, their mean free paths are distinctly smaller than those of photons. Therefore, surface sensitivity is obtained by taking electrons or ions as primary and/or secondary probes. For the methods discussed here this is the case for SEM, TEM, and STM. However, SFM is only sensitive to the surface, because interaction forces are repulsive on a short distance scale, and therefore prevent the massive cantilever (the probe) entering at least hard surfaces (but compare section 8.4). In many cases one is interested in ascertaining depth profiles. This can be achieved by sputtering (ion etching) the surface with ions, and removing the sample layer by layer while measuring. Because the sputter process changes the sample structure, depth profiles are not combined with microscopic techniques and will therefore not be covered here. Requirements of the techniques Techniques utilizing probes with a small mean free path are surface sensitive on the one hand, but need high or even ultra high vacuum conditions in the measurement chamber, because the probes have to travel from the source to the sample, and from there to the detector without scattering with gas particles. SEM

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8.5 Finite tips cannot enter small crevices on a sample surface.

and TEM cannot therefore be used if the samples have to be studied under atmospheric or liquid conditions. (For the exception of the environmental (E-) SEM compare section 8.3.) Charged particles often produce surface charging, therefore either conductive samples are required, samples have to be made conductive (e.g. by metal evaporation), or charge compensation (if possible) has to be applied. Optical methods and SFM do not suffer from these restrictions. STM needs conductive samples, but can be applied in air and in liquid because the electrons tunnel between the tip and sample surface at a very small distance, and there is no scattering possible. All scanning probe techniques are operated with thin solid beams which come into mechanical contact with the sample surface. Because these beams have a finite diameter, they cannot enter very narrow crevices (Fig. 8.5). Very rough samples are therefore not easily imaged.

8.2

Electron microscopies

In electron microscopes, electrons are used as primary probes. Due to the small mean free paths of charged particles in air, the microscope has to be held under high vacuum conditions. In the sample, a variety of different interactions between the probe electrons and the sample atoms can take place: · transmission of primary electrons; · secondary electron emission; · Auger-electron emission, a special type of secondary electron emission which involves three electrons. Firstly, an electron hole of high binding energy (produced by the incident electron beam) is filled by an electron of lower binding energy to reduce the overall energy of the system. The energy of this process is either emitted as X-rays (characteristic X-ray emission), or used to emit a third electron of low binding energy, the Auger electron. The energy of

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· · ·

·

209

this electron is characteristic for the element, hence a chemical analysis can be made. electron backscattering; absorption of electrons; characteristic X-ray emission (see above), also characteristic for a certain element. Differences with Auger electrons are that X-ray emission is more efficient for higher atomic numbers and that the information with X-rays stems from 1 m depth whereas Auger electrons show an information depth of a few 10 nm. emission of light.

One big disadvantage of electrons as a probe is charging of the sample during electron bombardment. In principle, only conducting samples can be imaged. In a special case non-conducting samples can be imaged if the number of the absorbed electrons is equal to the number of secondary electrons leaving the sample. In all other cases, the sample has to be coated by metal, or a conducting replica has to be made. Both methods may give artefacts. Furthermore, the samples have to be stable under electron irradiation. There are two approaches to electron-beam use in microscopy. The first is the transmission electron microscope realized by Ernst Ruska in 1933.6,7 The TEM can be compared with a transmitted-light-microscope: electrons pass through a thin sample8 and can be detected at a fluorescence screen. Because of the strong absorption and scattering processes which take place, if free electrons interact with matter, only very thin samples in the range of 10 nm to 100 nm thickness can be passed by electrons. Sample preparation is therefore difficult and timeconsuming and can produce artefacts. As in light microscopy there are different mapping modes:9 1.

2.

3.

Bright-field-image: the image is made of electrons which directly pass through the sample. The image plane of the objective is mapped by the projector lens. The contrast is provided by the loss of electrons by scattering and absorption on their transfer through the sample. This depends on the density of the sample and the atomic number of the sample components. Diffraction image: this also uses electrons which pass directly through the sample, but in this mode the focal plane of the objective is imaged by the projector lens. By this one gets a diffraction image which yields information on the crystal structure of the sample in a similar way to an X-ray-Lauediffraction image. Dark-field-image: electrons which scatter inside a sample are imaged on the screen, and the electrons which pass through the sample in a direct way are filtered out.

Additionally, energy-filtering transmission electron microscopy11 or TEM with attached energy dispersive X-ray analysis (EDX) is used which gives the elemental composition.

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The second approach to the use of electrons as a probe for imaging is the scanning electron microscope (SEM),12 which was invented by Knoll and v. Ardenne13 in the 1930s. The first realization of an image on a surface was described by Knoll und Thiele.14 This microscope scans a very well focused electron beam over the sample surface. The electrons which leave the sample surface are accelerated by a voltage to the detector, for example an EverhartThornley-Detector,9 which transforms the electrons into an electrical signal. This is used to control intensity in a cathode-ray-oscillograph, while the same signal that scans the electron beam in SEM is used as a deflection signal in the oscillograph. So an image of the electrons leaving the sample is obtained. Resolution in SEM is estimated by the area where the electrons leave the sample. The dimensions of this area depend on the spot radius of the electron beam on the sample surface and the scattering processes inside the sample. In SEM there are several kinds of secondary electrons which can be used for image formation: 1.

2. 3. 4. 5.

Secondary electrons which are produced by primary electrons near the surface with enough energy to pass through the surface barrier. These produce the highest resolution, down to 1 nm. Secondary electrons which are produced by backscattered electrons. These produce the lowest resolution, around 100 nm. Auger electrons, which have to be detected with a special energy resolving detector (e.g., a cylindrical mirror analyzer), and give a resolution of 10 nm. Secondary electrons which are produced from backscattered electrons out of the pole shoes of the last lens. Secondary electrons produced by primary electrons which impact the aperture and produce SEs there.

The last two cases only increase noise without any further information on the sample. Backscattered electrons are too fast to reach the secondary electron detector which is placed off-axis. They can, however, also be used for imaging with a different (in-axis) detector. Backscattered electrons give high material contrast because their number depends on nuclear charge and hence atomic number. The resolution obtained is around 100 nm. The resolution of SEM depends on how well the electron beam is focused on the sample. Typical resolutions are of the order of several nm. The depth of focus can reach up to 1 mm, if compared to only 100±1000 nm in a light microscope. A review of spatially resolved spectroscopies in electron microscopy for the study of nanostructures of different metals, semiconductors, and biomaterials is given in ref. 15. To overcome the limitation of a vacuum for biological samples and/or of charging of non-conducting samples, a so-called environmental SEM (E-SEM) was invented. In this a higher partial pressure can be held in the sample chamber (up to 10-4 bar) which is differentially pumped against the rest of the

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microscope. The ionization of the background gas helps to avoid charging. Also, the higher partial pressure of, e.g., water, keeps the biological structure almost in its native state, making this method a valuable tool for imaging biomaterials.

8.3

Scanning probe microscopies

Scanning probe techniques are a valuable tool since their invention in the 1980s.16,17 In these experimental methods, distance dependent interactions like tunneling current, force, or light transmission between a sharp needle (`tip' or `probe') in close proximity to a surface (`sample') are utilized to produce an image of the sample. Two principal measurement modes were implemented: (i) to maintain a constant height of the probe above the sample while measuring the interaction change reflecting surface topography (`constant height mode'), and (ii) to maintain constant interaction while adjusting the height with the feedback signal reflecting surface topography (`constant interaction mode') (Fig. 8.6). Both modes offer the possibility of characterizing surfaces down to the atomic scale in a great variety of environments from ultra high vacuum to aqueous solutions. It is also possible to characterize time dependent reactions like crystallization and corrosion processes, as this can be done continuously and hence on-line. To scan a probe over a surface in the desired way, while reacting to the topography of the sample, positioning tools with spatial resolution in the 0.1 nm-regime are necessary. The latter is fulfilled by piezoelectric actuators

8.6 Different modes of operating scanning probe methods (a) constant height mode, (b) constant interaction mode.

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made of ceramics like PZT (lead zirconate titanate) or PMN (lead magnesium niobate), which in different directions can be extended by less than the size of one crystal unit cell. With a suitable detection unit to measure small interaction changes connected to a distance feedback circuit and digital data representation, an image of the surface can be portrayed. Experimental adaptation and extensions based on the interaction mechanisms originally used for imaging purposes lead the way to monitoring more complex features than topography, as described below. A readable overview covering many scanning probe techniques is available.18,19 In scanning tunneling microscopy (STM), the current from electrons tunneling between a conductive wire (preferably heavy metals like tungsten, platinum or iridium) with an atomically sharp tip and a (semi-)conductive surface across vacuum is measured (Fig. 8.7). The tunneling current I decays exponentially with increasing distance s between tip and surface: p ÿ p
eff 8:2 V exp ÿk eff s I s with V as the applied (Bias-)voltage, eff the effective work function, and k a constant. I additionally depends on the local densities of electronic states of the tunneling partners: X f …Ei †‰1 ÿ f …Ef †ŠR2fi …Ef ÿ Ei † 8:3 I fi

Ei,f is the energy of the initial (i) or final (f) state of the tunneling event, f(E) the Fermi function giving the probability of occupation of the energy E, and Rfi the tunneling matrix element. The delta-function stands for the energy conservation of the process. Because of its strong distance behavior, only the atom at the very end of the tip and the nearest surface atom are involved within the tunneling event. A slight change in distance, according to progression along rows of

8.7 Schematic drawing of the STM principle.

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surface atoms, alters the tunneling current in a measurable manner, so that true atomic resolution can be achieved. Experimentally, a tunneling current is obtained by securing a tip to surface distance of less than a few tenths of a nanometer under constant bias voltage. The sign of the bias determines the direction of the tunneling current, which means that unoccupied electronic surface states are probed with occupied electronic tip states or vice versa. Scanning the surface row by row either at constant height or constant current (see above) reveals surface topography. Additionally, STM can be applied to probe electronic structures. Modulating the tip-surface distance and measuring the change of the tunneling current at constant bias allows extraction of the local workfunction U of the sample. Modulating the bias and measuring corresponding current changes at a constant tip-surface distance allows extraction of the electronic states around the Fermi level of the sample (STS, scanning tunneling spectroscopy). This can lead to chemical information about the surface, but for more detailed information, electronic core levels have to be sensed which is not possible with STM (see ref. 20 and references therein). Scanning electrochemical microscopy (SECM) measures highly localized electrochemical currents associated with charge transfer reactions on metallic sample surfaces under a liquid environment.21,22 In macroscopic measurements, it can be compared with cyclic voltammetry. The reactions can occur in a fourelectrode electrochemical cell under bipotentiostatic control. There are two pathways for image production. Electron tunneling and electrochemical reactions via an electrolyte bridge occur according to the applied voltage; this can be used for the detection of localized electrochemical reactions at surfaces. It can also be used for microstructures of biomaterials like titanium.23 In addition, SECM is capable of probing the kinetics of solution reactions, adsorption phenomena and monitoring heterogeneous electron transfer kinetics associated with processes at conducting surfaces.24 The invention of scanning force microscopy (SFM) was a breakthrough for these techniques, as it became possible to image non-conducting substrates with a resolution of 0.2 nm laterally and 0.001 nm vertically. It does not require a specimen to be metal coated or stained. Non-invasive imaging can be performed on surfaces in their native states, and under near-physiological conditions. It has proven to be particularly successful for imaging biological samples such as proteins, nucleic acids and whole cells. By scanning, dynamic processes can be imaged such as erosion, hydration, physicochemical changes, and adsorption at interfaces. Therefore SFM is currently the SPM technique with the widest applicability for biomaterial research. In SFM a small tip attached to a micro-beam (cantilever) is scanned across the surface of a specimen, and deflected by topographic features (Fig. 8.8(a)). The force of interaction may be repulsive or attractive, giving rise to different modes of operation. Moving the cantilever from the interaction free zone far

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8.8 (a) Schematic drawing of the SFM principle, (b) Cantilever deflection versus sample-z-position-curve on a stiff surface monitored via SFM.

above the surface it snaps into contact (Fig. 8.8(b)) due to attractive force between the tip and sample, which can be described in a simple way by the Lennard Jones potential. The piezo pushes the tip further towards the sample and the positive repulsive force reaches a maximum. As the piezo is retracted the repulsive force is reduced, and the force changes sign. If the bending force of the cantilever becomes greater than the attractive force towards the surface, the tip loses contact. The tip can be held in the repulsive regimen of the Lennard Jones potential or oscillated in an attractive or repulsive regime resulting in different interactions.25,26 These differences are important, as biomolecules are deformed by applying a load of some nN as present in contact mode, i.e., in the repulsive regime.27 Deflection is usually monitored by a laser beam that is reflected and detected with a four split photodiode. This signal is used to maintain a constant force via a feedback loop, and to monitor the height data. SFM can be operated in a variety of modes that can provide different information about the sample. Usually the z-deflection is monitored, and interpreted according to the parameters under investigation. This is mainly done via monitoring of the zpiezo voltage. This can cause ambiguities, as piezo crystals exhibit hysteresis. This is overcome by monitoring the distance separately via inductive or fiberoptic sensors.28 For dynamic modes29 the application of an additional oscillation to the cantilever by a piezo crystal has to be performed. Magnetically driven cantilevers are also used for actuation.30 The quality of SFM data is essentially determined by the cantilever and the tip, influencing the resolution of topography and force measurements. Micromechanical properties of the cantilever, and the shape and chemical composition of the tip, which comes into direct

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contact with the sample, are essential. High aspect ratios and small tip radii are desirable for imaging steep slopes and deep crevices. Depending on the mode of operation, different parameters have to be optimized, which can be realized according to the methods described in the following. Silicon and silicon nitride cantilever fabrication based on photolithographic techniques are well established. Metal based (Ni) cantilevers31 and cantilevers made of piezoelectric material (lead zirconate titanate) are also produced for independent actuation and sensing.32 Cantilevers of various shapes (e.g. rectangular or V-shaped) and dimensions (usually 100 to 200 m in length) are available. New approaches to further miniaturization of the system have been made on microfabricated aluminum probes with length scales of 9 m.33±36 Coatings (Cr-Au, Pt, Al, TiN, W2C, TiO2, Co, Ni, Fe, Au with biological coating) can be applied to the cantilever to modify it against corrosion in the liquid phase, or for different applications when conductive, magnetic, or biological properties are necessary. The resonance frequency of the cantilever ranges from several kHz to several hundred kHz. Cantilevers with high resonance frequencies are used in dynamic mode as the tip oscillates at several hundred kHz above the surface. The stiffness of the cantilever is defined by the force constant k, and ranges between 0.01 to 100 N/m for the vertical deflection. Soft (k < 0.1 N/m) cantilevers are used in contact mode to minimize disturbance of the sample. Rigid cantilevers with a force constant larger than 1 N/m are used in non-contact or dynamic modes since they exhibit high resonant frequencies and small oscillation amplitudes of several nanometers. The force constant for lateral twisting can be determined for friction measurements,37 and the movement of the cantilever has been modeled accordingly with finite element analysis.38 The mechanical behavior and the determination of the spring constant is well described in the literature.39±46 The tip itself can consist of different materials, or is coated according to the application. For some applications, a hard surface is necessary and a diamondlike coating (DLC) is applied.31 The production of DLC coatings has been described,47 but other functionalities48±50 can be applied. Thus cantilever tips can be modified by chemical51 and biochemical52 functionalization. With silanization,53,54 or via thiols, self-assembled monolayers (SAMs) allow further modification.50 Proteins and bacteria55 can also be attached to the tip. Additional materials which vary the shape of the tip can be deposited, such as polystyrene, borosilicate and silica spheres, C60 molecules,56 carbon-nanotubes in general57±60 or single wall nanotubes with diameters of pI, which was attributed to electrostatic repulsion due to surface charge on the proteins and increased spreading (conformational changes) because of reduced protein stability.193 QCM-D measurements at multiple harmonics were used to investigate a number of biomolecules in the adsorbed state.190,194 The QCM data were further supported by D2O substitution and ellipsometry measurements, which were used to quantify the amount of coupled water detected by the QCM. The quartz crystal sensor surface can be coated, in principle, with any biomolecule and utilized for the development of biofunctional patterned surfaces. This was done by extending lipid- and protein-based immobilization strategies to surfaces patterned with alternating areas of Au and SiO2. After a number of sequential immobilization steps, all of which were followed either in parallel or simultaneously using QCM-D, Au was functionalized by DNA, while SiO2 was covered by an inert phospholipid bilayer. Since the phospholipid bilayer is not only inert towards protein adsorption, but also towards lipid vesicle adsorption, this pattern allows controlled specific immobilization of DNAtagged lipid vesicles, which may or may not carry protein functionality.190,194 QCM has been used to study the fouling of metals exposed to protein solutions and the efficiency of surfactants in cleaning up metal surfaces.195 The QCM surface coated with gold or chromium was exposed to 3 wt% solutions of pure -lactoglobulin and a commercial skimmed milk powder at neutral pH, and before and after heat treatment at 80 ëC. The piezoelectric sensor was also used to evaluate the cleaning action of surfactant Tween 20 and enzyme trypsin. The mass load changes on the QCRS were interpreted in terms of the hydrodynamic thickness of the adsorbed films and led to the conclusion that QCM is a sensitive and convenient technique for monitoring detergent efficiency.195

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A recurrent problem in QCM applications is the reproducibility of the experimental results. Lipids and proteins can adsorb irreversibly and nonspecifically onto the quartz crystal surface making its regeneration for subsequent re-use a significant challenge.77,196 It has been suggested that proteases are used for surface cleaning from proteins or to use disposable sensors.196 QCM coated with a pH-sensitive amphoteric polymer has been used to monitor cell growth.197 The sensor responded to the biochemical products of cellular metabolism rather to cells, and it did not require immobilized biological receptors to monitor cell growth and metabolism. Microbial biofilm formation on metallic surfaces has been studied by a combination of EQCM with gravimetric and optical methods.198 The use of flow-cell EQCM allowed real-time monitoring of biofilm formation and resulting open circuit potential UR. Measuring UR was particularly useful for assessing different methods for surface cleaning from microbial biofilms. It was shown that surface treatment with 10% H2O2 completely restored the original UR value, whereas disinfection with 70% ethanol did not affect it.198 The quartz crystal resonant sensor was capable of measuring the adhesion of Staphylococcus epidermidis to fibronectin-coated surfaces. Changes in resonant frequency were recorded and showed a linear relationship with the logarithm of cell suspension concentrations ranging from 1  102 to 1  106 cfu/ml.199 A non-invasive, real-time QCM method for the monitoring of cellular integration within commercial collagen-based dermal replacement scaffolds was reported.188 An unexpectedly high degree of acoustic energy transfer through heavily hydrated thick films (up to 0.5 mm) of collagen/glycosaminoglycan scaffold materials intimately associated with a quartz crystal sensor allowed for quantitative resonant frequency measurements on application of fibroblast cell suspensions to the material. Changes in resonant frequency and energy dissipation were commensurate with cellular interaction with the gel. Many marine organisms attach to solid surfaces with an extraordinarily strong adhesion. QCM-D was used to study adhesives from two marine organisms, the common blue mussel Mytilus edulis, and the brown algae Laminaria digitata. The cross-linking ability of adhesive byssal proteins from M. edulis (Mefp-1) is based on the presence of the reactive amino acid 3,4dihydroxyphenylalanine (DOPA), which can be oxidatively crosslinked by the enzyme tyrosinase. The brown algae L. digitata contains polyphenolic biopolymers which can be oxidatively crosslinked with bromoperoxidase (BPO). Chemical cross-linking of both proteins can be done with NaIO4. Adsorption of Mefp-1 and PP was measured by a decrease in resonant frequency and increase in dissipation. Addition of NaIO4 resulted in a significant decrease in frequency and viscoelasticity of the adsorbed layers due to crosslinking, which was also confirmed by ellipsometry.200 Monitoring cell growth or adhesion using QCM seems to offer a significant advantage over biological methods in terms of complexity of the method and

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duration of the analysis, but like protein adsorption, it is doubtful that there is a simple linear correlation between the frequency shift and the mass of the cells attached to the crystal surface.157,201 The frequency shift calculated for a monolayer of cells attached to a 5 MHz QCRS using the Sauerbrey equation (13.1) was at least ten times higher than the experimentally observed
f .201 The most likely explanation of this discrepancy is that the cell layer has visco-elastic properties and energy dissipation should be taken into account. Even very thin (a few nm) biofilms dissipate a significant amount of energy owing to the QCM oscillation.202 Various mechanisms contribute to this energy dissipation. Three main contributions suggested by Rodahl et al.202 are: (i) viscoelasticity and porosity of the biofilm that is strained during oscillation, (ii) liquid entrapped between the quartz crystal surface and the biofilm layer, which can move between or in and out of the pores due to the deformation of the film and (iii) the load from the bulk liquid which increases the strain of the film. These mechanisms constitute an effective visco-elastic load hence, biofilms cannot be considered to be rigidly coupled to QCM oscillation and data interpretation of QCM measurements of biofilms requires more sophisticated models. A comprehensive model for `adhered cells-quartz surface' system has been suggested.203 It considers three layers, namely a rigid layer of the extracellular matrix, a visco-elastic layer of cells and a water layer between them. The QCM-D method used to study cell attachment and spreading over the polystyrene coated quartz crystal surface showed correlation between the degree of dissipation and amount of cells attached.204 An unusual QCM set-up allowed very sensitive detection of bacteriophage.205 By immobilizing ligands that interacted specifically with phage coat proteins on the QCRS surface, selectivity of phage adsorption was achieved; then the amplitude of the surface oscillation was then increased to `shake off' the adsorbed phage. The authors claimed that the quantification range spanned over at least five orders of magnitude and the method was sensitive enough to detect as few as twenty phages. Such sensitivity makes this method attractive for virus detection. Not only can QCM be used to detect cells, but living cells immobilized on the QCR sensor can also be used to study effect of chemical substances on their behaviour. The QCM technique measuring
f and
R detected microtubule (MT) alterations caused by the drug nocodazole in the immobilized endothelial cells (EC).206 Depolymerization of microtubules was dose dependent in the concentration range of nocodazole between 0.11 and 15 M. By relating QCM data to the microscopic examination of endothelial cells, it has been shown that different regions on
f ÿ t and
R ÿ t curves correspond to different events that change mass and viscoelasticity of the EC layer.207 These events reflect different stages of cell-surface interactions such as contacting of the surface, adhersion and spreading. In the examples quoted above the quartz crystal resonance sensor technique was either used as only a mass change detector, or as a mass and dissipation

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detector. It is also possible to design a QCR biosensor that exploits only the viscosity-density change of the liquid medium. In an experimental set-up the concentration of a biomolecule is measured in the solution by carrying out a sol-gel transformation which significantly influences the viscosity of a medium.208,209 The latter is detected by a quartz TSM sensor. If there is no film formation on the quartz crystal surface, the frequency and resistance changes of the sensor depend only on the liquid density and viscosity. This approach was realized to measure the concentration of fibrinogen in a clotting assay and endotoxin in the limulus ameobocyte lysate LAL test which is also based on the clotting of the LAL substrate and the endotoxin. It has been suggested that the Kanazawa±Gordon equation (13.4) and eqn 13.10 are applicable in these systems:210 r  `  ` 13:10
R ˆ 8K 2 C0 f0 q q where K 2 is the electromechanical coupling factor of the quartz and C0 is static quartz capacitance. A polystyrene spin-coated QCR with a fundamental frequency of 10 MHz was used to record frequency and resistance changes during the LAL-endotoxin clotting. Different approaches to produce a calibration plot were examined and it was found that the time to reach a certain resistance shift gave a linear graph `time vs logarithm of endotoxin concentration' at low endotoxin levels. At higher endotoxin concentrations, the
R or
f at a certain time gave a better correlation; the combination of both methods allowed fast measurement of endotoxin concentration in the range 100 pg/L to 2 g/L.210 QCM immunosensors Immunosensors are biosensors in which the immunochemical reaction is coupled to a transducer.211,212 The fundamental basis of all immunosensors is the specificity of the molecular recognition of antigens (Ag) by antibodies (Ab) to form a stable complex. This is similar to immunoassay methodology. Immunosensors can be categorized by the detection principle used. The main types are electrochemical, optical and microgravimetric immunosensors. In contrast to the enzyme-linked immunosorbent assay (ELISA), which is most commonly used to detect Ag-Ab interactions, modern transducer technology enables the label-free detection and quantification of the immune complex. Further potential advantages of immunosensors vs conventional immunoassays include lower cost and a shorter time of analysis. The analysis of trace substances in environmental science, and pharmaceutical and food industries is a challenge, since many of these applications demand a continuous or a nearcontinuous monitoring. The use of immunosensing sequences in these applications is a possibility. Similarly, continuous monitoring of certain analytes

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can be used in clinical diagnostics. The future role of this technique in the laboratory as well as for bedside testing will become even more important as the clinical laboratory is faced with increasing pressure to contain costs.211 Shons et al.213 first applied antigen coated piezoelectric crystals to immunoassay. The first piezoelectric immunosensors lacked reproducibility and stability when immersed in a liquid, so a measurement technique known as `dip and dry' was used.213±215 According to this protocol, the QCR sensor (QCRS) with immobilized antibody was dipped into a solution containing antigen, rinsed, dried and
f was compared with that of the sensor before dipping. Using this technique, herbicide atrazine, Candida albicans and IgM (human) were detected. In one case atrazine was detected in the range from 0.001 to 1 ppb, which is a remarkable sensitivity.216 Nevertheless, dip-and-dry technique suffers from intrinsic problems ± it is slow and cannot be used in the continuous regime. More commonly QCM for immunosensing is used in the continuous mode in liquid phase.23,217 The key issue in achieving stability and reproducibility of analysis is the choice of the immobilization method of the immune reactant on the quartz crystal surface. As with biosensor fabrication, immobilization can be done in a variety of ways. Physical adsorption of antibodies is an attractive method due to its simplicity and it can be universally used to coat any surface.218 This is based on on irreversible adsorption of high molecular mass protein molecules.219 Electrical impedance analysis has been used to study Ab-Ag interaction on a polystyrene-coated quartz crystal surface.220 The motional resistance, R increases during both the Ab immobilization and the subsequent Ab-Ag binding indicating more power dissipation in the system. Resonant frequency, f decreases in both cases reflecting increased mass loading.
R in the Ab-Ag reaction is considerably larger than that in antibody immobilization process while
f are similar in these two processes. The ratio
R/
f therefore increases three-fold for Ab-Ag interaction vs. Ab immobilization.220 Physical adsorption is a statistical process driven by van der Waals forces, and adsorbed molecules have an irregular orientation; moreover, the adsorbed proteins undergo conformational changes that affect their properties such as bioactivity.221 Using AFM, it was shown that non-covalent attachment of antibodies ± sheep or mouse IgG ± to the gold coated mica resulted in large aggregated structures comprising randomly oriented antibodies.218 Not surprisingly, such a random orientation resulted in a lower efficiency of AgAb binding as measured by QCM. Another complication is the Vroman effect, that is, displacement of one set of adsorbed proteins by other proteins.222 These significant drawbacks of immobilization by physical adsorption make it a less popular technique for analysis in the liquid phase than alternative methods. Problems with low stability and reproducibility of sensors with physically adsorbed antibodies led to development of alternative approaches for binding of

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immune reactants to the QCR sensor surface. The Au electrode surface does not have surface sites for a strong attachment of biomolecules; to create these, several techniques have been used. They include the formation of: (i) a polymer layer with functional groups; (ii) a molecularly imprinted polymer layer; (iii) a self-assembled monolayer with functional group, (iv) an adsorbed layer of molecules with high affinity towards the immune reactant. In the latter case, formation of the interacting couples `protein A-immunoglobulin G', `streptavidin-biotin' or `avidin-biotin' is most commonly used. The initial step in the immobilization procedure is the preparation of bare QCRS electrode surface. In most cases the electrode is made from gold ± an inert material whose properties do not change in time. Its surface, however, may get contaminated due to inevitable adsorption of volatile organics and water upon storage.223 Although cleaning of the electrode surface has not always been used prior to immobilization,224±228 it has been shown that the frequency of cleaned QCM stabilizes faster that that of non-cleaned QCM,229 and most researchers carry out cleaning of the electrode surface before its modification.215,217,230±233 Pre-treatment of the QCR sensor in a glow discharge plasma chamber is used as a method of its surface cleaning.234 The majority of protocols for surface cleaning include treatment with a strong alkali solution (usually NaOH) followed by rinsing and treatment with a strong acid (HCl).235,236 Alternatively, the gold electrode surface is cleaned using `Piranha' solution (one part of 30% H2O2 in three parts of concentrated H2SO4).218,237,238 Polymer coating of QCM electrodes is commonly used to derivatize their surface, through which biomolecules are attached.239 Polymers can be immobilized by dip-coating,231,233,235 in which case the polymer is usually spread from a non-aqueous solution on the electrode and then air-dried. Polymer coating by adsorption from aqueous solution is based on hydrophobic interaction between the electrode surface and polymer.229 Polyethylenimine is often used for this purpose, as it has reactive amine groups through which immune reactant molecules can be attached. Koning and GraÈtzel231 elaborated a sensor (coating with polyethylenimine followed by glutaraldehyde, prior to the attachment of antiglycophorin A antibody) for human erythrocytes. Coating by chemical attachment of ( -aminopropyl)triethoxysilane ± (silanization) is also used to create reactive amine groups on the surface.233 For the immobilization of biological materials on the electrode surface, electropolymerization has some advantages, because this process is easier to control and films have a more uniform thickness and chemical composition and are mechanically and chemically stable.240,241 Glow discharge plasma polymerization has been used to make thin polymer films from ethylenediamine and allylamine, which also contain amine functionality.234,242,243 Further immobilization of antibodies, antigens or other biomolecules is accomplished using adsorption or a chemical coupling agent such as glutaraldehyde.229,234,235 Immobilization of antibodies

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via aldehyde groups of glutaraldehyde on a quartz crystal precoated with PEI produces a layer of randomly oriented antibodies.235 Considering the high costs of QCM sensors as the main obstacle towards commercialization, Su suggested use of silver electrodes instead of gold.244 Silver is less inert and less stable than gold and a polystyrene coating was used to protect the silver electrode from undesired oxidation and immobilize biomolecules to make an immunosensor.196,245 Lin et al.246 have argued that molecularly imprinted polymer layers (MIP) formed on the quartz crystal surface in the presence of analyte have affinity and selectivity comparable to that of natural receptors, and they are easier to prepare than Ab-based sensors. The absence of biomolecules on a QCRS ensures its higher physico-chemical, mechanical and thermal stability. A mixture of 3-dimethylaminopropyl methacrylamide with different acrylate crosslinkers and albumin was copolymerized and spin-coated on a Au-coated quartz crystal surface. Albumin was extracted from the MIP with 20% methanol in water. Thus formed, the MIPQCM showed higher selectivity towards albumin vs other proteins of a lower molecular mass using model mixtures and blood serum. The MIP sensor also had faster response times and greater adsorption capacity towards analyte. The covalent coupling of the Fab' fragment of antibodies via its free thiol groups to a monolayer of phosphatidylcholine and cholesterol on QCM was used to create an immunosensor with high antigen binding efficiency.247 Protein A, a cell surface receptor produced by Staphylococcus aureus, has been widely used for the quartz crystal surface coating in immunosensor design. This protein has high natural affinity with the Fc region of immunoglobulins, particularly IgG molecules.248 Despite the absence of chemical interaction between protein A and IgG, the dissociation constant KD for the protein A-IgG complex is very low; it varies between 10ÿ9 M and 10ÿ6 M assuming pseudofirst-order of this process.249 For comparison, highly specific Ab-Ag interactions form complexes with KD values around 10ÿ9 M.250 Such a high affinity between protein A and immunoglobulin provides stable binding of IgG to QCRS surface coated with protein A. One protein A molecule can bind at least two IgG molecules.251,252 Protein A is a single polypeptide chain of molecular weight 42,000.251 If physically adsorbed on a QCRS surface without losing its ability to bind IgG molecules and without blocking the active sites of the antibodies for binding target analyte.218,253 Although protein A adsorption is of a physical nature, driven by van der Waals forces, it is irreversible like any other protein adsorption, and the gold-protein A complex is highly stable.217 Guilbault et al.215 utilized polyclonal antibody attached to the gold electrodes pre-coated by protein A. A big advantage of using protein A coating is that antibodies can be directly attached to it and no additional surface activation is required,218,235 whereas immobilization of IgG molecules on other polymer films requires use of a coupling reagent that activates surface functional groups.

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Attachment of antibodies to the surface via protein A has the advantage of preferential orientation of the immobilized molecules, as their Fab domains are likely to be exposed to the external medium.217,231,254 Use of protein A for immobilization of antibodies produced better stability and reproducibility of results in QCM detection of Salmonella typhimurium compared to the QCM sensor designed by covalent attachment of antibodies to the PEI film via glutaraldehyde activation. The sensitivity of both sensors was comparable.235 Like antigen-antibody interaction, streptavidin (avidin)-biotin interaction is highly specific and irreversible, with KD values reported as low as 10ÿ16 M.241 It has been used for the immobilization of biomolecules in a QCM immunosensor. Dupont-Filliard et al. coupled biomolecules to the QCM surface coated with polypyrrole-biotin film through an intermediate avidin layer. The procedure involved three steps: (i) electrosynthesis of the polypyrrole-biotin film at an electrode surface; (ii) surface immobilization of the avidin layer via the biotin entities and (iii) anchoring of the biotin-labeled biomolecules on the polypyrrole biotin/avidin layer to create a bioactive surface. The authors described several advantages of using this technique, namely, (i) avoiding the use of chemical reagents that may affect biomolecules, (ii) high precision of immobilization of biomolecules only at the surface of the film, which leads to a high sensitivity, (iii) versatility of the technique, as a wide variety of commercially available biotin conjugates is available. A DNA sensor was prepared by immobilization of biotinylated DNA probes creating the sensing layer, polypyrrole-biotin/avidin/ DNA probe. The sensor was regenerated by destroying the biotin/avidin complex using a detergent solution via the solubilization of the avidin layer. Removal of the avidin layer immobilized on a polypyrrole-biotin film did not alter the functionality of the latter. After the initial 15±20% loss of activity, regenerated biosensors exhibited a stable and reproducible response up to the 10th regeneration. Use of self-assembled monolayers in the QCM immunosensor design is based on the high affinity of thiolated compounds with a gold surface.255 The SAM technique is one of the simplest ways to produce a well-ordered layer suitable for further modification with antibodies. Based on the immobilization of antibodies onto a SAM of cystamine with sulfo-succinylmidyl 4-(pmaleimidophenyl)butyrate (sulfo-SMPB) as linker, a piezoelectric immunosensor was fabricated for sensing C. trachomatis with a detection limit of ~260 ng/ml.256 Thiols are strongly attached to gold surface via chemisorption rather than physical adsorption forces. Hydroxy-, amino- and carboxylterminated thiols are commonly used for covalent immobilization of bioligands to SAM. A great variety of terminal functionalities in monolayers offers high flexibility in biosensor design using SAMs. Zhao et al. 257 coupled hydroxyl and amino-terminated SAMs to the 50 phosphate-end of DNA, while carboxyl-terminated ones were coupled to the 30 hydroxy end of the nucleic acid using a water soluble carbodiimide. X-ray photoelectron spectroscopy (XPS) and cyclic voltammetry (CV) data proved

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that hydroxyl-terminated SAMs were more efficient for the covalent attachment of DNA than the other two types of SAM. A multi-layer QCM for Chlamydia trachomatis was designed by assembling a primary layer of cystamine on the gold electrode; the cystamine monolayer was further modified with sulfosuccinylimidyl 4-(p-maleimidophenyl)butyrate and the fragmented F(ab')(2) anti-mouse IgG Ab, to which the anti-C. trachomatis LPS-Ab was attached. The latter molecule was used for Chlamydia trachomatis detection. The sensor sensitivity for this bacterium was 260 ng mLÿ1 in urine and it showed long-term stability upon storage at 4 ëC. 256 Similarly, a piezoelectric immunosensor based on a self-assembled monolayer of cystamine has been developed for the determination of Schistosoma japonicum Ab in rabbit serum. Immobilization of the Schistosoma japonicum Ag to the positively charged SAM was achieved electrostatically using a negatively charged polystyrene sulphonate layer. It was shown that Ag immobilized by this procedure had higher immunological activity and binding efficiency than Ag immobilized with glutaraldehyde. The immunosensor had satisfactory sensitivity and detection limit.258 To increase the mass load on a QCM immunosensor, an increase in surface area for the Au electrode was suggested by coating it with porous gold. A 16fold increase of surface area was achieved, and this led to a 11.4-fold increase in thiol adsorption and a 3.3-fold increase in protein adsorption, thus amplifying the sensitivity of the original flat non-porous device.259 A piezoelectric immunosensor was developed for the rapid detection of Escherichia coli O157:H7. Antibodies were immobilized onto a monolayer of 16-mercaptohexadecanoic acid self-assembled on an Au electrode at an AT-cut quartz crystal. Covalent attachment of Ab to the carboxylic groups of the SAM was carried out with N-hydroxysuccinimide ester as a reactive intermediate. The immunosensor could detect the target bacteria in a range of 103±108 CFU/ml within 30±50 min. The proposed sensor was comparable to Protein A-based piezoelectric immunosensor in terms of the amount of immobilized antibodies and detection sensitivity.260 Similarly, a batch-type QCM system for detecting chloramphenicol (CAP) was developed. Anti-CAT Ab were covalently attached to the carboxylic groups of SAM of 3-mercaptopropionic acid via activation with water-soluble carbodiimide (EDC) and N-hydroxysulphosuccinimide. The antibody-immobilized sensor showed 10±50-fold enhanced sensitivity in comparison with the uncoated sensor or coated only with 3-mercaptopropionic acid. Repeated use of the sensor up to eight times was possible after 1 min regeneration with 0.1 M NaOH.236 These results proved that the QCM sensing has potential as a rapid screening method for low molecular mass compounds such as environmental endocrine disruptors, which usually require timeconsuming, complex and expensive labelling analytical techniques. The SAM technique can be used to immobilize biomolecules by labelling target bioligands with thiol groups. Thus, biomolecules become capable of selfassemblying on a QCM surface. Using this approach, thiolated DNA and

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oligonucleotides were used to produce DNA hybridization sensors.261 Immobilization of thiolated Ab produced variable results, perhaps due to the oxidation of ±SH groups.218 Molecules with disulphide links were suggested, e.g. DNA±(CH2)3±S±S±(CH2)3±OH, to minimize the undesirable side reaction of ±SH oxidation. The disulphide link can then be cleaved by reduction with dithiotreitol prior to immobilization. Thiolated anti-Salmonella antibodies to Salmonella typhimurium were directly attached to the Au electrode of a QCM sensor from a solution of Ab, and a heterobifunctional thiolation cross-linker, sulphosuccinimidyl 6-[3-(2pyridyldithio)propionamido]hexanoate.262,263 The immunosensor could detect 9.9  105 to 1.8  108 cells/ml within 30±90 min. For repeated use of the immunosensor, the efficiency of various regeneration reagents in removing adsorbed Salmonella cells and maintaining sensor sensitivity were compared. The best results were achieved with the use of 1.2 M NaOH. After the fifth assay, the decrease in frequency was only 24.4%, whereas treatment with 8 M urea solution led to the decrease in a frequency of 43.6% and treatment with 0.2 M glycineHCl, pH 2.8 resulted in a frequency shift by 63.2%. Rickert et al. also reported that regeneration of the QCM immunosensor with a 6 M urea solution. After this experiment, the sensitive layer was regenerated by destroying the antigen-antibody binding with a 6 M urea resulting in a 30% loss of sensitivity after three regeneration cycles.264 Several papers describe immunoassays in which the QCM sensor measures a frequency shift caused by agglutination of immunized latex microbeads. The method was named by its authors latex piezoelectric immunoassay, or LPEIA.265±268 A suspension of latex microbeads (0.2 m in diameter) with immobilized antibodies against the analyte reacts with antigen, as a result of the interaction, the latex particles agglutinate. An uncoated QCM registers an increase of mass as the oscillation frequency shift due to the precipitation/ adsorption of the agglutinated particles. LPEIA was used to measure clinically relevant fibrin degradation products (FDP) and C-reactive protein (CRP) in blood serum. Response of the QCM sensor to FDP in human serum was within 10 min and stabilized within 60 min. The frequency shift of the QCM sensor and the absorbance change at 570 nm of the photometric method showed good correlation. As the piezoelectric sensor is used uncoated, its regeneration can be carried out with piranha solution. The sensor was used at least three times without any loss of sensitivity and results were reproducible.269 A commercial prototype sensor for CRP detection has been designed. Similarly, changes in viscosity/density of solution resulting from an antigenantibody agglutination reaction were followed by the QCM technique.270 Using this approach, Staph. epidermidis was detected in the infected human serum samples. Use of QCM immunosensors is not limited to the detection of immune reactants or biomolecules. Environmental pollution resulting from dioxins is

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becoming huge problem in the world. A piezoelectric immunoassay was developed for the rapid detection of polychlorinated dibenzo-p-dioxins (PCDDs). The system uses a competitive inhibition enzyme immunoassay (EIA) based on a mouse monoclonal Ab against 2,3,7,8-tetrachlorodibenzo-pdioxin (TCDD) and a conjugate of a dioxin-like competitor coupled to the enzyme horseradish peroxidase (HRP). The anti-dioxin Ab was immobilized on a 10 MHz AT-cut quartz crystal resonator modified with a SAM of dithiobis-Nsuccinimidyl propionate. TCDD at different concentrations in the range 0.001± 10 ng mLÿ1 was mixed with a constant amount of HRP-conjugated competitor and the frequency response due to the adsorption of the samples on the biosensor surface measured. The results showed that TCDD could be quantitatively detected in the concentration range 0.01±1.3 ng mLÿ1. The sensitivity and selectivity of the QCM immunosensor is comparable to EIA and ELISA methods in the detection of PCDDs. The developed QCM immunosensing system offers significant improvements in speed, sample throughout and cost for the qualitative and quantitative detection of PCDDs compared with GC-MS.271 Regeneration of the modified sensor interface by dissociating the bound analyte from the antibody-coated sensor indicates the reusability of this biosensor. In this study, both 20 mM glycineHCl buffer (pH 2.5) and 6 M urea were used to regenerate the QCM biosensor. The results indicate that the sensing interface still retained more than 86% activity after four regeneration cycles performed with glycine±HCl buffer (pH 2.5). The recovered activity using 6 M urea was 70% after four regeneration cycles. To detect the herbicide atrazine, monoclonal Ab against atrazine were immobilized on a quartz crystal sensor via Protein A covalently attached to the gold electrode surface activated with 3,3-dithiobis(propionic acid) Nhydroxysuccinimide ester. The MAb±Protein A complex was stabilized by cross-linking with dimethyl pimelimidate. The immunosensor thus developed was able to specifically respond to atrazine with a linear response of the frequency shift to atrazine in the concentration range 1±200 g/mL, however, sensor regeneration and re-use were a problem.272 Anti-PCB polyclonal sheep Ab were immobilized on quartz crystal sensor via protein A. The QCM sensor thus obtained was successfully used to detect 4,40 DCB (dichlorobiphenyl) and 2,4,40 -TCB (trichlorobiphenyl); the sensor detected TCB in the concentration range 0.2±2000 g/L and, remarkably, it worked in non-aqueous solutions as well.273 Cooper et al. suggested a novel approach for using QCM immunosensing techniques.274 Instead of measuring the frequency shift caused by the interaction between type 1 herpes simplex virus (HSV1) and specific antibodies covalently attached to the QCM surface, the authors monitored detachment of the virions from the surface by monotonously increasing the amplitude of oscillation of the piezoelectric sensor, while using it to detect the acoustic noise produced when the interactions were broken. The method termed rupture event scanning

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(REVS) is quantitative over at least six orders of magnitude of concentration range, and its sensitivity approaches detection of a single virus particle (0.5 femtogram) in phosphate buffer solution containing 1 mg/mL BSA and 50 virions in serum. Sensitivity of REVS is several orders of magnitude better than that of QCM or SPR. The REVS technique has been successfully used to detect E. coli, N. meningitidis, adenovirus type 5 and S. aureus. Scan time is as short as 1 min and no amplification is required for detection.

13.4 Combination of QCM with other techniques 13.4.1 QCM and SPR Surface plasmon resonance (SPR) is a highly sensitive method currently used for monitoring processes occurring at or near interfaces.275 It is an optical method based on total internal reflection from an interface between a material with higher refractive index such as metal film on glass and a medium with lower refractive index such as liquid. The generated evanescent wave moves away from the interface into the medium penetrating it to a distance of around 300± 600 nm. Analysis of the evanescent wave profile gives information about the adsorbed layer and allows measuring its thickness and amount of the substance adsorbed.276 SPR has been used as a gas phase sensor and as a method of monitoring biochemical and cellular interactions in a liquid.277±279 Like QCM, it is a non-invasive, non-labelling technique that provides kinetic data in real time. SPR is often combined and compared with QCM.238,280 QCM measurements are more sensitive to the environment, such as high ionic strength solutions, drifts in temperature or mechanical disturbance that cause changes in solution viscosity and density. With tight temperature control, SPR measurements have better reproducibility and reliability. Careful solution preparation and temperature control improve QCM performance. On the other hand, the QCM instrumental setup is relatively simple and fast. Both methods are easily capable of monitoring protein deposition with high sensitivity, however, the resonance angle shifts observed in SPR when studying bacterial suspensions are very low. In the same system QCM was capable of measuring bacterial concentration as low as 1000 cells/ml although the reliability of the method has to be improved.238 In another system studied, interaction of glucose with its specific enzyme glucose-6-phosphate dehydrogenase (GDH), SPR failed to register the enzyme-substrate binding, whereas the QCM flow system clearly showed strong interaction between GDH and glucose. Apparently change of the refractive index was too small in this system.185,237 Because QCM and SPR monitor changes in different parameters it is not surprising that for certain applications one technique is more sensitive than the other. An important advantage of QCM over SPR may be a relative simplicity of surface modification of the piezoelectric crystal surface. This approach was used

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to study complement activation by different biomaterial surfaces such as titanium, polystyrene or poly(urethane urea).281 The sensitivity of both methods was comparable although it depended on the analytical cell design.

13.4.2 QCM and other techniques EQCM is the most obvious example of QCM coupled with another technique in the same instrument and its applications are discussed in other paragraphs of this chapter. In most cases QCM and other methods are used complementarily. Hess et al. used the scanning electrochemical microscope (SECM) in combination with EQCM to investigate standing acoustic waves over the quartz crystal.282 The adsorption kinetics of three human blood proteins ± human serum albumin, fibrinogen and haemoglobin from model solutions onto glass and quartz coated with 12 nm thick titanium dioxide film was measured using three different experimental techniques: optical waveguide light spectroscopy (OWLS), ellipsometry (ELM) and QCM-D.190 All three techniques proved to be suitable to study kinetics of protein-surface interactions in situ. They were also consistent in recording the adsorption kinetics trends and the results were used to quantify the adsorbed mass. Using different protocols of surface treatment, it was shown that protein adsoption was largely irreversible, and so was the adsorption of antibodies against these proteins. The two optical techniques produced comparable results in most cases, which could be converted into adsorbed protein (`dry') mass. Data obtained with QCM-D, on the other hand, differed significantly in terms of the adsorbed mass, being 1.75 to 3.2 times higher. The higher mass value calculated from
f includes both protein mass and water associated with the protein layer. Analysis of the energy dissipation in the adsorbed layer and its magnitude in relation to
f provides information about the mechanical and structural properties, such as viscoelasticity and shear modulus, of the adsorbed film and the surrounding bulk liquid environment.190,202,283 The combination of optical techniques and QCM revealed much more information than any of these methods used separately. By measuring the damping of the amplitude of the crystal vibration after the alternating electric field has been turned off, one can qualitatively determine how much energy the adsorbed layer can dissipate. Since the signal from rigid adlayers is only weakly damped, it can be said to have a small dissipation factor. In other words, there are relatively few degrees of freedom, into which energy can be lost. Conversely, soft adlayers have many degrees of motional freedom and a correspondingly large dissipation factor. The amplitude of the signal from a crystal with such a layer adsorbed on it decays very rapidly. The QCM-D instrument serves as a complement to optical waveguide lightmode spectroscopy. Both techniques are capable of low-background, real time, in situ measurements of ng/cm2 levels of adsorption of unlabeled samples from aqueous solution. However, they provide different information since OWLS can

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only detect the mass of the adsorbed molecule itself, whereas QCM is also sensitive to entrapped or closely coupled solvent molecules such as water. In addition, visco-elastic information inherent in the value of the dissipation factor has no equivalent in OWLS. The QCM-D has the further advantage that it does not require transparent substrates, allowing examination of substrates such as gold in addition to the metal oxides used in OWLS experiments.190,202,283 It has been pointed out that QCM is one of the few methods that allows precise measurement of the quantity of an ultrathin layer of an element deposited onto a solid surface.284 Other methods such as Auger or X-ray photoelectron spectroscopies provide only relative and semi-quantitative information, whereas nuclear or ion-scattering spectroscopies are limited to certain combinations of the adsorbate and substrate. Unlike QCM, all these methods require expensive equipment and complicated experimental conditions. Results obtained with inductively-coupled-plasma mass spectroscopy, used to determine the coverage of an ultrathin layer of antimony on gold, were consistent with QCM measurements on the same sample.284 Coupling of the tapping-mode atomic force microscope (AFM) and QCM has an advantage of obtaining simultaneous information at nano- and sub-nano(AFM) and micro- (QCM) levels.285,286 The AFM/QCM coupled technique proved useful for investigating a wide range of processes, from protein adsorption on metallic surfaces285 to metal electrodeposition.286 Copper and silver deposition on gold produces rigid films, whereas protein adsorption is dominated by viscous interactions. Respectively, the frequency variations of the overtone number n evolve as
fn =n for rigid films, while in the case of viscous interactions in which the thickness of the liquid interacting with the resonator is a function of the oscillation frequency, the frequency variations of the overtone number n evolve as
fn =n0:5 . Measuring overtones thus provides useful quantitative and qualitative information on the nature of the adsorbed film. A coupled quartz crystal microbalance/heat conduction calorimeter (QCM/ HCC) has been designed for very sensitive and simultaneous mass and heat flow measurements of thin films interaction with a solvent vapor in a controlled gas phase.287

13.5 Acoustic/piezoelectric sensors Acoustic/piezoelectric sensors are based on a piezoelectric substrate through which acoustic waves propagate. QCM operates with a bulk acoustic wave (BAW) and its mass sensitivity is limited by the crystal thickness.288 The surface acoustic wave (SAW) devices can operate using different modes of wave propagation.22,289±291 Rayleigh SAW sensors are suitable for gas media but not for liquids in which they suffer unacceptably high energy dissipation and loss of sensitivity because of the surface-normal particle displacement.67,292 In contrast, shear-horizontal polarized surface acoustic wave modes (SH-SAW) or

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plate modes (SH-APM) are not significantly affected by immersion in liquid because the particle displacement in these modes is horizontal ± parallel to the sensor surface. Currently they are becoming a subject of intensive studies as biosensors.293,294 SH sensors are less sensitive than the Rayleigh SAW sensors, so to increase their sensitivity the waves are propagated through a waveguiding layer.295,296 First SH-APM sensors showed poor sensitivity that has not been significantly improved.297 Among such devices those using Love waves have recently attracted particular attention due to their sensitivity.* Love waves are elastic SH waves that are generated in the substrate and guided by a thin film deposited on it.298 Their propagation is sensitive to mass loading and viscoelastic effects of the medium. Usually two interdigital transducers (IDT) ± transmitting and receiving ± are used. Love-wave devices are based on ST-cut quartz crystal as the substrate with a guiding layer made from a polymer such as PMMA,294 deposited or etched SiO2,299,300 or a bi-layer structure comprising silica and PMMA.292 The PMMA layer is more sensitive than the silica layer but the latter is chemically inert, so their combination in a bi-layer has some advantage over individual layers.292 The mass sensitivity, stability and temperature coefficient of frequency of the bi-layer Love-mode quartz acoustic sensor is better than those of PMMA- or SiO2-coated devices but the insertion loss in water remains significant. It has been estimated that Love-wave devices should be sensitive enough to detect a uniform monolayer of most substances.301 A typical value of operating frequency of Love wave devices is about 100 MHz, much larger than typical values of QCM devices, near 10 MHz.302 High sensitivity of 500 pg/ml was reported for IgG adsorption to a Love mode mass sensor with the guiding ZnO layer (4 m thickness) on ST-cut quartz substrate. To increase selectivity of adsorption, an additional film of Au (20 nm thick)/Cr (10 nm thick) was deposited on the zinc oxide layer.303 ST-cut quartz has a low dielectric constant compared to water, hence being suitable for less polar liquids, its small piezoelectric coupling coefficient makes it less efficient in aqueous solutions.67 An alternative substrate for SH-SAW sensor is based on a 36ë rotated Y-cut X-propagating LiTaO 3 . 304 It has high dielectric constant and high electromechanical coupling coefficient. With a protective layer of SiO2 (1 m thickness) and an additional gold top layer (50 nm thickness) a SAW immunosensor was designed. It was coated with protein A to which rabbit anti-mouse IgG F(ab)2 fragment (antibody) was immobilized and subsequently the sensor was used to detect the goat anti-rabbit IgG (antigen). The sensor showed expected selectivity but further improvement of the sensor design would be required to improve its sensitivity.67 An in vitro assay to monitor the specific binding of an ultrasound contrast agent to a receptor immobilized on Love-wave biosensor has been developed.305 An antibody was coupled to the sensor via the * They are named after the English mathematician AEH Love, who described them in 1911.

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carboxylic groups of the monolayer of 11-mercaptoundecanoic acid selfassembled on the surface. Love-wave immunosensors with immobilized AM13-PVIII antibodies against a major phage coat protein PVIII have been successfully tested in detection of M13 bacteriophage.302 The antibodies were immobilized on the surface using dithiobissuccinimidyl-propionate. For the lowest bacteriophage concentration tested (1.6  109 pfu mlÿ1) the relative frequency shift was between ÿ14 ppm at 10 min and ÿ69 ppm at 2 h, which was better than ÿ7 ppm reported for a QCM based sensor.306 The response time was relatively slow for high concentrations of bacteriophages ÿ2h, but better than is usually required for routine biological analysis. In commercially available SAW transducers interdigital structures are usually made from aluminium or chromium by photolithographic technique. These electrodes corrode even in pH neutral buffer solutions, so that surface coating of SAW sensors is also necessary to protect the electrodes against corrosion.307 Protective polymeric films can also be used for covalent immobilization of biomolecules.308 It has been suggested that for covalent immobilization of biomolecules to SAW sensor surface, milder conditions and reagents should be used than for fabricating QCM biosensors because commercially available lithium tantalate SAW devices contain aluminium structures. Examples of immobilization methods suitable for engineering SAW biosensors include BrCN activation of the polyimide film,309 silanization of silica coating310 and photoimmobilization.308 Photoimmobilization requires derivatization of protein molecules with photoreactive groups. To produce sensing layers, proteins were modified at the lysine group with 3-fluoromethyl-3-(m-isothiocyanophenyl)diazirine (TRIMID). Upon light activation, covalent coupling of the protein molecule to the polyimide coating is achieved.308 Similarly, a dextran layer was attached to the surface. Dextran has a versatile chemistry allowing further covalent attachment of biomolecules.311,312 BrCN activation technique and immobilization of anti-mouse IgG (antibody) led to the design of a mass immunosensor with sensitivity of 58 Hz/pg by mouse IgG (antigen) and 51 Hz/ pg for the immobilized anti-glucose oxidase (antibody) by glucose oxidase (antigen)309 but the whole immobilization process took a week to complete and reproducibility of the sensor was unsatisfactory. Aluminium structures corroded in the course of the immobilization. On the contrary, photoimmobilization of proteins proved to be reproducible but it produced lower sensitivity of detection and still was quite complex a procedure.308

13.6 Future development of piezoelectric sensors There are general trends in the development and analytical applications of acoustic/piezoelectric sensor techniques in liquid media, which should further improve their performance:

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· improvement of sensors and whole electronic circuits; · consideration of energy dissipation, impedance, conductivity, surface nonuniformity and roughness, stress and compressional waves using comprehensive theoretical and numerical models and development of relevant software; · application of piezoelectric sensors in combination with other techniques such as spectroscopy, microscopy and chromatography; · use of sensor arrays for analysis of complex biosystems; · use of new piezoelectric materials such as GaPO4 may become more widely applied; · development of new setups possessing greater sensitivity than conventional QCM method; · miniaturization of sensors and cells; · increase in working frequency of QCM crystals to 30 MHz and higher; · improved stability and reproducibility of other acoustic/piezoelectric sensors; · direct measurements of adsorption of small molecules specifically interacting with surface functionalities; · development of new experimental setups for real-time and on-line use of piezoelectric sensors; · new more specific methods of surface modification and immobilization of bioreceptors; · immunoassays in organic solvents convenient for investigations of hydrophobic molecules; · investigation of bioaffinity interactions on the interfaces using advanced methods (impedance scans, REVS, etc.); · development of new biosensors using a variety of biofunctionalities with protein, e.g., RNA, DNA, lipids, saccharides, enzymes.

13.7 Thermal gravimetry A thermal analysis (TA) technique enables efficient measurement of changes of physical parameters such as mass (TG, DTG), temperature (DTA), enthalpy (DSC), together with various physical properties determined through mechanical (TMA), optical, magnetic, electrical, acoustic, emanation and other observations. These measurements take place during heating of materials using a controlled temperature program. Evolved gas analysis (EGA) may be based on chemical measurements, depending on the method of detection used. Consequently, an essential subsequent interpretation step is usually required to relate these physical TA measurements to each chemical change that is of interest.313 Available instruments permit a rapid and efficient completion of thermal experiments, obtaining quantitative data that are suitable for measuring weight loss upon heating, reaction enthalpy, specific heat, reaction temperature, melting point and reaction rate. Thermal changes, including chemical reactions,

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are often more complex than recognized, so that kinetic data and their interpretation may be incorrect. In measuring rates of thermal changes, the following factors may influence kinetic characteristics: reaction reversibility, melting, reactant self-cooling/heating, multiplicity of rate processes (concurrent and/or consecutive), and others.313 The term TKA (ThermoKinetic Analysis) describes any research in which physical measurements are usually carried out to determine the course and/or extent of the changes occurring in a sample held at constant or programmed temperature environment. It allows one to elucidate the reaction mechanism (in the usual chemical meaning of this term) and identify parameters that determine absolute and relative levels of reactant reactivity and control reaction rates. These objectives are generally approached through experimental investigations which attempt to find out what chemical reactions have occurred (stoichiometry), how fast (kinetics) and obtain any other information capable of elucidating all relevant aspects of the changes identified such as textures from microscopy, structures and topotaxy from crystallography, and bonding from spectra.313,314 Enthalpy changes can contain composite contributions, for example, where there are two (or more) concurrent chemical reactions and/or a phase change, such as progressive melting and/or reactant sublimation. Some types of change in the reacting system are not detected by particular TA methods, for example, TG does not record enthalpy changes due to a reaction, melting or other crystallographic transformations. To overcome method-specific limitations, complementary TA techniques may be used; TG and/or DTG are often accompanied by DSC and EGA with chemical and/or physical detection methods. Another limitation is that TA techniques, in contrast with microscopic observations, do not give direct evidence of the changing (geometric) dispositions within the reactant particles. Reaction geometry is often inferred indirectly through the interpretation of reaction rates by heterogeneous, interface advance, kinetic models.313±315 There have been many publications concerned with general and specific aspects of the experimental methods and theory of TKA, 316 together with specialist surveys and theoretical discussions.313±315 These citations refer to an extensive, now predominantly older literature, which has been recently revised.316,317 Current concepts of crystal chemistry, that form the theoretical foundation for TKA kinetic analyses, are incapable of constituting a comprehensive culture for continuing investigations of the variety of thermal changes that occur on heating reactants in condensed phases.313 Thermal gravimetry is a destructive method, which has lower sensitivity than QCM and has limited applications in investigation of interfacial phenomena in liquid phase.

13.8 Non-QCM adsorption methods Measuring of the adsorbed amount of gas and recording the adsorption isotherms are widely used for analysis of the surface structure of materials.318,319

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These measurements can be performed by gravimetric or volumetric methods. Conventional methods for the adsorption of gases and vapours such as N2, Ar, Xe, CO2, H2O and C6H6, and mercury porosimetry are characterized by much lower sensitivity (by several orders) than the QCM technique in respect to the studied weight per surface area unit. However, the adsorption isotherms of probe gases measured by standard methods allow one to investigate the structural characteristics of hard porous materials, including their porosity, pore size distribution, specific surface area, distribution of the adsorption potential and fractal dimension. Application of a combination of methods for the comprehensive characterization of surfaces, interfacial phenomena and adsorption of gases and solutes is fruitful. It should be noted that non-QCM gravimetric methods typically work in the gas phase and at reduced pressure, and this limitation is a barrier to investigations of biosystems in native liquid media. For many porous biomaterials used in tissue engineering which are soft hydrogels rather than rigid structures, these methods are hardly applicable, and other non-gravimetric methods are required.320

13.9 Dynamic contact angle measurements Contact angle between a sessile liquid droplet and a flat solid surface as a wetting phenomenon was originally described for non-porous and non-absorbing surfaces. Static contact angle  is defined geometrically as the angle formed by a liquid at the three-phase boundary where a liquid, gas and solid intersect. If the angle is less than 90ë the liquid is said to wet the solid. If it is greater than 90ë it is said to be non-wetting. A zero contact angle represents complete wetting.321±324 The static contact angle is measured directly by examining the shape of a sessile droplet on flat surface ± goniometry method. A sessile droplet on a tilted flat surface has two different angles ± advanced and receded. These angles are also obtained directly by examining geometry of the droplet. If the three-phase (liquid/solid/vapour) boundary is in actual motion, the angles produced are called dynamic contact angles (DCA) and are referred to as `advancing' and `receding' angles. They can be measured indirectly using microgravimetry ± tensiometry method. Tensiometry involves measuring the forces of interaction between a solid and a test liquid. The difference between the maximum (advanced/advancing) and minimum (receded/receding) contact angle values is called the contact angle hysteresis.325±328 A great deal of research has been dedicated to the analysis of the nature and significance of hysteresis. It has been used to characterize surface heterogeneity, roughness and mobility. Briefly, on non-homogeneous surfaces there will be domains which present barriers to the motion of the contact line. On chemically heterogeneous surfaces, these domains represent areas with different contact angles to the surrounding surface. For example when wetting with water occurs, hydrophobic domains will pin the motion of the

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contact line as the liquid advances thus increasing the contact angles. When water recedes, the hydrophilic domains will hold back the draining motion of the contact line thus decreasing the contact angles. From this analysis it can be seen that, when testing with water, advancing angles will be sensitive to hydrophobic domains and receding angles will characterize hydrophilic domains on the surface. For geometrically rough surfaces, the actual microscopic variations of the slope on the surface create the barriers, which pin the motion of the contact line and alter the macroscopic contact angles. Contact angle can also be considered in terms of the thermodynamics of the materials involved. This analysis involves the interfacial free energies between the three phases and is given by:

lv cos  ˆ sv ÿ sl

13:11

where lv , sv and sl refer to the interfacial energies of the liquid/vapour, solid/ vapour and solid/liquid interfaces. In the case of porous solids, powders and fabrics, the Washburn method is commonly used. If the forces of interaction, geometry of the solid and surface tension of the liquid are known, the contact angle may be calculated. The sample of the solid to be tested is hung on the balance and tared. The liquid is then raised to contact the solid. When the solid contacts the liquid, a sharp change of interaction forces is detected and a tensiometer registers this elevation as zero depth of immersion. As the solid is pushed into the liquid, Ftotal is measured by the microbalance: Ftotal ˆ wetting force ‡ weight of sample buoyancy force

13:12

Modern DCA equipment has software that tares the weight of the probe and removes the effects of the buoyancy force by extrapolating the graph back to zero depth of immersion. The remaining component is the wetting force defined as: Wetting force ˆ LV P cos 

13:13

where LV is the liquid surface tension, and P is the perimeter of the probe. This contact angle, which is obtained from data generated as the probe advances into the liquid, is the advancing contact angle. The sample is immersed to a set depth and the process is reversed. As the probe retreats from the liquid, the data collected is used to calculate the receding contact angle. The use of tensiometry for measurement of contact angle has several advantages over conventional goniometry. At any point on the immersion graph, all points along the perimeter of the solid at that depth contribute to the force measurement recorded. Thus, the force used to calculate  at any given depth of immersion is already an averaged value. One may calculate an averaged value for the entire length of the sample or average any part of the immersion graph data to assay changes in contact angle along the length of the sample. This technique allows one to analyze

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contact angles produced from wetting over an entire range of velocities from static to rapid wetting. The solid is brought into contact with the testing liquid and the mass of liquid absorbed into the solid is measured as a function of time. The amount absorbed is a function of the viscosity, density and surface tension of the liquid, the material constant of the solid, and the contact angle of the interaction. If the viscosity, density and surface tension of the liquid are known, the material constant and contact angle can be determined. The primary focus of contact angle studies is in assessing the wetting characteristics of solid/liquid interactions. The work of adhesion is defined as the work required to separate liquid and solid phases, or the negative free energy associated with the adhesion of solid and liquid phases. It is used to express the strength of the interaction between the two phases and is expressed by the Young-Dupre equation as: Wa ˆ …1 ‡ cos †

13:14

Work of cohesion is defined as the work required to separate a liquid into two parts, it is a measure of the strength of molecular interactions within the liquid. It is expressed as: Wc ˆ 2

13:15

Work of spreading is the negative free energy associated with spreading liquid over solid surface. Also referred to as a spreading coefficient it is given as: Ws ˆ …cos  ÿ 1†

13:16

Wetting tension as a measure of force/length is defined as: t ˆ Fw =P ˆ LV cos 

13:17

This value, wetting force normalized for length, also represents the product of the cosine of the contact angle and the surface tension. It allows for a characterization of the strength of the wetting interaction without separate measurement of surface tension, and is most helpful in characterizing multicomponent systems, where surface tension at an interface may not be equal to the equilibrium surface tension. Measurement of contact angles yield data, which reflect the thermodynamics of a liquid/solid interaction. Using a series of homologous liquids of differing surface tensions a graph of cos  vs is produced; it is found that the data form a line which approaches cos  ˆ 1 at a given value of . This is the maximum surface tension of a liquid that can completely wet the solid. This value, called the critical surface tension, is used to characterize a solid surface. Another way to characterize a solid surface is by calculating surface free energy, also referred to as the solid surface tension. This approach involves testing the solid against a series of well characterized wetting liquids. The liquids used must be characterized so that the polar and dispersive components of their surface tensions are known. The relevant equation derived by Owens and Wendt is:

Microgravimetry

l …1 ‡ cos †=… ld †1=2 ˆ … sp †1=2 ‰… lp †1=2 =… ld †1=2 Š ‡ … sd †1=2

365 13:18

where l is the liquid surface tension and s is the solid surface tension, or free energy. The subscripts d and p refer to the dispersive and polar components of each parameter. The form of the equation is linear of the type y ˆ mx ‡ b. One can plot … lp †1=2 =… ld †1=2 vs l …1 ‡ cos †=… ld †1=2 . The slope gives … sp †1=2 and the y-intercept gives … sd †1=2 . The total free surface energy is merely the sum of its two component forces. In the case of liquid adsorption by a porous solid, a weight correction should be considered   x ÿ xZDOI ; 13:19 F ˆ p L cos q ‡ A
g…x ÿ XZDOI † ‡
m 1 ÿ xmax ÿ xZDOI

where p is the object perimeter, L is the surface tension of a liquid, g is the acceleration due to gravity, x is the position of the plate and xZDOI is the position of the plate at zero depth of immersion,
m is the change in mass of a sample after complete withdrawal from the liquid. A method based on QCM for measuring sessile contact angles and surface energies of liquid-air and liquid-liquid interfaces has been described.329 This method involves measurement of the frequency change accompanying the introduction of a small liquid droplet to the center of a vibrating quartz resonator, which comprises an AT-cut quartz crystal sandwiched between two gold electrodes. If the density and viscosity of the liquid are known, the contact angle between the droplet and the gold substrate surface can be determined directly. QCM measurements of contact angles formed between aqueous droplets and gold surfaces modified with various organosulfur monolayers having different surface energies agree with sessile contact angles determined by optical goniometry. Furthermore, the QCM method can be used to measure the contact angle formed between an aqueous droplet and the QCM surface when both are submerged under an immiscible solvent such as hexadecane. In this case, the frequency change relies on the differences in the densities and viscosities of the water droplet and the fluid displaced by the droplet at the surface. The dependence of contact angle on the concentration of surfactant in a aqueous droplet provides for determination of the critical micelle concentration for aqueous phases in contact with air or an immiscible organic fluid. These measurements can be performed under conditions where contact angles cannot be measured readily, such as in the presence of opaque media or in the case of two liquids having similar refractive indexes.329 It is tempting to relate the surface energy of a solid to its biocompatibility, but our understanding of this relationship is poor.330±332 Even although sometimes it is possible to obtain a correlation between the surface energy of a biomaterial and its performance, as found for PMMA coated with a biomimetic phosphorylcholine coating,333,334 this relationship is far too complex to draw definitive conclusions.

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13.10 Conclusion Modern and most prospective developments of micro-, nano-, and femtogravimetry are based on the crystal sensor technology, which gives fruitful in-depth results for many complex systems, such as cells, biomacromolecules, polymers, thin films, drugs, and interfacial phenomena. Other gravimetric methods, direct (electronic balances, thermogravimetry and isothermal adsorption) or indirect (dynamic contact angle), are characterized by a much lower sensitivity than the crystal sensor techniques. There is an important trend in the applications of the crystal sensor techniques by combining with other methods such as AFM, SEM, TEM, SPR, XPS and specroscopy, that provides detailed information about complex interfacial phenomena, which will have both great theoretical and practical significance for the future.

13.11 Acknowledgements This work has been supported by DTI (UK) project MPP4.2. V. M. G.'s visit to the University of Brighton (UK) was supported by NATO (CLG Grant No. 979845).

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284. Narine S S, Hughes R and Slavin A J, `The use of inductively coupled plasma mass spectrometry to provide an absolute measurement of surface coverage, and comparison with the quartz crystal microbalance', Appl Surf Sci, 1999 137(1±4) 204±6. 285. Choi K H, Friedt J M, Frederix F, Campitelli A and Borghs G, `Simultaneous atomic force microscope and quartz crystal microbalance measurements: Investigation of human plasma fibrinogen adsorption', Appl Phys Lett, 2002 81(7) 1335±7. 286. Friedt J-M, Choi K H, Francis L and Campitelli A, `Simultaneous atomic force microscope and quartz crystal microbalance measurements: Interactions and displacement field of a quartz crystal microbalance', Jpn J Appl Phys Part I, 2002 41 (6A) 3974±7. 287. Smith A L and Shirazi H M, `Quartz microbalance microcalorimetry a new method for studying polymer-solvent thermodynamics', J Thermal Anal Calorim, 2000 59(1±2) 171±86. 288. Du J, Harding G L, Collings A F and Dencher P R, `An experimental study of Love-wave acoustic sensors operating in liquids', Sensor Actuat A ± Phys, 1997 60(1±3) 54±61. 289. Kalantar-Zadeh K, Trinchi A, Wlodarski W and Holland A, `A novel Love mode device based on a ZnO/ST-cut quartz crystal structure for sensing applications', Sensor Actuat A ± Phys, 2002 100(2±3) 135±43. 290. Cheeke J D N and Wang Z, `Acoustic wave gas sensors', Sensor Actuat B ± Chem, 1999 59(2±3) 146±53. 291. Zhang C, Caron J J and Vetelino J F, `The Bleustein-Gulyaev wave for liquid sensing applications', Sensor Actuat B ± Chem, 2001 76(1±3) 64±8. 292. Du J and Harding G L, `A multilayer structure for Love-mode acoustic sensors', Sensor Actuat A ± Phys, 1998 65(2±3) 152±9. 293. Andle J C, Vetelino J F, MW Lade M W and McAllister D J, `An acoustic plate mode biosensor', Sensor Actuat B ± Chem, 1992 8(2) 191±8. 294. Gizeli E, Stevenson A C, Goddard N J and Lowe C R, `A novel Love-plate acoustic sensor utilizing polymer overlayers', IEEE T Ultrason Ferr, 1992 39(5) 657±9. 295. Thompson D F and Auld B A, `Surface transverse wave propagation under metal strip gratings', Proc IEEE Ultrasonics Symp Williamsburg, VA, USA, 261±6, 1986. 296. DeÂjous C, Savart M, RebieÁre D and Pistre J, `A shear-horizontal acoustic plate mode (SH-APM) sensor for biological media', Sensor Actuat B ± Chem, 1995 27(1±3), 452±6. 297. Dahint R, Ros Seigel R, Harder P, Grunze M and Josse F, `Detection of non specific protein adsorption at artificial surfaces by the use of acoustic plate mode sensors', Sensor Actuat B ± Chem, 1996 36 (1±3), 497±505. 298. Gulyaev Y V, `Review of shear surface acoustic waves in solids', IEEE T Ultrason Ferr, 1998 45(4), 935±8. 299. Kovacs G, Vellekoop M J, Haueis R, Lubking G W and Venema A, `A Love sensor for (bio)chemical sensing in liquids', Sens Actuators A, 1994 43(1±3) 38±43. 300. Jakoby B and Vellekoop M J, `Analysis and optimisation of Love-wave liquid sensors', IEEE T Ultrason Ferr, 1998 45(5) 1293±302. 301. Du J, Harding G L, Ogilvy J A, Dencher P R and Lake M, `A study of Love-wave acoustic sensors', Sensor Actuat A ± Phys, 1996 56(3) 211±19. 302. Tamarin O, Comeau S, DeÂjous C, Moynet D, RebieÁre D, Bezian J and Pistre J,

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Part III Surface interaction and in-vitro studies

14 Interaction between biomaterials and cell tissues Y I W A S A K I and N N A K A B A Y A S H I , Tokyo Medical and Dental University, Japan

14.1 Introduction In developing biomedical materials, we are concerned with their function, durability, and biocompatibility.1 Particularly, the surface properties of materials are directly related with this concern. Durability, particularly in a biological environment, is less well understood. Still, the tests we need to evaluate durability are clear. The important question in biocompatibility is how the device or material `transduces' its structural makeup to direct or influence the response of proteins, cells, and the organisms that are to relate to it. For devices and materials that do not leach undesirable substances in sufficient quantities to influence cells and tissues, this transduction occurs through the surface structure ± the body `recognizes' the surface structure and responds to it. For this reason we must understand the surfaces of biomaterials and the interfaces between biosubstances and biomaterials. In this chapter we describe the surface phenomena of biomaterials. First, the general recognition of a living organism to artificial materials and biocompatibility, which is a characteristic common to all materials, are discussed. Second, typical methods for surface analyses are introduced. Third, a design for `non-biofouling' surfaces, which are useful for materials that are exposed to blood or for diagnosis, is described. Finally, bioconnective materials based on the concept of a reliable system for connection to tooth substrates are discussed.

14.2 Surface properties of biomedical materials 14.2.1 Bioreactions on biomaterial surfaces2 A large amount of effort goes into the design, synthesis, and fabrication of biomaterials and related devices to ensure that they have appropriate mechanical properties, durability, and functionality. These functionalities of biomaterials and medical devices are primarily derived from the bulk structure of the material. In contrast, biological responses to materials are influenced by their

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14.1 Classification of interfacial biocompatibility.

surface chemistry and structure. Therefore, surface design and characterization are quite important in the development of biomaterials. Biocompatibility is a complementary definition for the required surface properties of biomaterials. `Biocompatibility is the ability of a material to perform with an appropriate host response in a specific application' (Williams, 1987).3 Thus, the meaning of biocompatibility depends on the specific application of the materials used. The roles of biomaterials can be divided into two categories, as shown in Fig. 14.1.4 One is `non-stimulative' and the other one is `bioconnective'. Because non-stimulative properties of materials complement non-activation, antithrombogenicity and tissue non-invasion are listed properties. These properties are necessary for materials that are exposed to blood. Conversely, bioconnective materials are required for adhesives for hard and soft tissue, sealants, and tissue engineering applications. Although these categorized materials will probably be used in completely different situations, the same surface properties, which can control interaction with the biocomponents, are required. Non-stimulative materials must reduce nonspecific protein adsorption and cell adhesion, socalled `biofouling.' Conversely, bioconnective materials should accumulate specific proteins and cells on their surfaces. However, it is usually difficult to control the interaction at the interface between artificial materials and a living organism. When these materials are exposed to a vital environment, non-specific protein adsorption and cell adhesion quickly occur. This biofouling reduces the functionality of the materials and also induces unexpected bioreactions such as thrombus formation, immune responses, complement activations, capsulations, etc.5,6 These reactions are important for maintaining vitality, however, they are serious problems for medical treatments that employ artificial materials. Most implanted medical devices serve their recipients well for extended periods by alleviating the conditions for which they were implanted, improving the quality of life and, with some device types, enhancing survival. However, some implants and extracorporeal devices ultimately develop complications that lead to device failure and thereby may cause harm to the patient or even death.

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Table 14.1 Typical bioreaction on artificial materials Protein adsorption Protein retention Lipid adsorption Bacteria adhesion Hemolysis Platelet adhesion Platelet activation Expression of new gene

Macrophage adhesion Phagocytosis Macrophage release Neutrophil activation Biodegradation Angiogenesis Cell spreading Fibrous encapsulation

All implants interact to some extent with the tissue environment in which they are placed. The biomaterial-tissue interactions encountered most frequently are summarized in Table 14.1. The important difficulty associated with medical devices is largely based on biomaterial-tissue interactions that include both the effects of the implant on the host tissue and the effects of the host on the implant. Several of the most important interactions in clinical and experimental implants and medical devices include inflammation and the `foreign body reaction,' immune response, systemic toxicity, thrombosis, device-related infection, and tumorigenesis. These interactions make up aberrations of physiological processes that function as common host defense mechanisms and are induced by non-specific biofouling. Thus, it is critical to prepare novel polymers that can control interactions with biocomponents at their interfaces.

14.2.2 Surface design of biomaterials1 To improve biointerfaces, the surface properties of biomaterials such as chemistry, wettability, topology, charge density, and bioactivity have been modified. Table 14.2 summarizes methods of surface modification and their efficiency. Coating is the most convenient method of surface modification. However, the affinity of a coating material to a substrate must be considered to reduce elution of the coating polymer. Graft polymerization is one of the familiar methods used to modify surface properties; the resulting modified surface is relatively stable because the graft polymers are connected with the substrate by covalent bonding. Graft polymerization can be achieved by several treatment methods such as application of chemical reagents, plasma-, corona-, photo-irradiation, etc. In particular, the corona discharge treatment is very convenient because it can be performed in an ambient atmosphere without the use of chemical reagents. Polymer blending is also an interesting method of surface modification in both research and technology. However, obtaining a homogeneous mixture of two or more polymers is difficult due to their high molecular weights. Moreover, the hydrophobic moiety shows a tendency to accumulate at the surface because

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Table 14.2 Surface modifications with polymer materials Modification methods Treatment

Effects

Chemical reaction

Oxygen functional groups Oxygen functional groups Sulfonic acid Thiol group Defluorine, oxygen functional groups Amination Oxygen functional groups

Chromic acid micture Chloric acid-sulfuric acid Sulfuric acid Sulfur Sodium Amine Sodium hypochlorite

Modifications of the original surface

Electron beam Corona discharge Plasma etching

Grafting

Radiation grafting Photografting Plasma, corona

Depends (cross-linking, degradation) Oxygen functional groups Oxygen functional groups Various polymers

Noncovalent coatings Solvent coating Layer by layer adsorption Vapor deposition

Various polymers

Biological modification

Peptide, protein

Covalent bonding Intermolecular force

of the decrease in the surface free energy. In general, protein adsorption easily occurs on hydrophobic surfaces due to hydrophobic interaction. Although the stability of surface modification with blending would be better than that with coating, the disadvantage of blending is that surface properties are too difficult to control. An inter-penetrating polymer network (IPN) is defined as a network composed of two chemically independent cross-linked polymers. IPN technologies contribute properties from each of two independent polymers to a material. Monomers from the surface of a substrate penetrate and cross-link with the monomers in the subsurface of the substrate, giving a density-graded modification layer. This is a suitable method for surface modification because the reduction of the mechanical properties of the substrate can be reduced. There have been numerous studies for the fixing of bioactive molecules to materials to give their surfaces biological functions. For non-stimulative materials, particularly, blood antithrombogenic surfaces, heparin and thrombomoduline were fixed. In contrast, cell adhesive peptides and proteins were also immobilized on conventional polymer surfaces to improve affinity to cells. As expected, some of these bioactive molecules worked on the surface for short periods. However, their activity often declined meaning that they are inadequate for use in long-term medical treatment. In addition, to obtain bioactive molecules with high efficiency, the surface designs of the substrates

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are important because nonspecific biofouling and interaction with the substrate also reduce functionality.

14.3 Surface analyses of biomedical materials1 14.3.1 Introduction Surface analysis is quite important for understanding the properties of biomaterials. Of course, the methods of surface analyses of materials are constantly being improved. Table 14.3 shows the characteristics of many common methods of surface analysis, including the depth of analysis and the spatial resolution. For surface analysis, we must consider several points. The surface structure of the material is often mobile. The movement of the atoms and molecules near the surface in response to the outside environment is often highly significant. In response to a hydrophobic environment, more hydrophobic components may migrate to the surface of a material. In some methods of surface analysis, materials are kept under a condition of an ultra high vacuum. This condition is completely different from a physiological environment. Characterization of a Table 14.3 Typical surface analysis methods for biomaterials Method

Principle

Depth analyzed

Area analyzed

Contact angles

Liquid wetting of surfaces is used to estimate the surface energy

0.3±2 nm

1 mm

XPS: X-ray photoX-rays cause the emission of electron spectroscopy electrons of characteristic energy

1±25 nm 10±150 m

SIMS: secondary ion mass spectroscopy

Ion bombardment leads to the 1 nm±1 m emission of surface secondary ion

10 nm

FTIR-ATR: Fourier transform infra-red spectroscopy, attenuated total reflectance

IR radiation is adsorbed in exciting molecular vibrations

1±5 m

10 m

SPM: scanning probe Measurement of the quantum microscopy tunneling current (STM) or van der Waals repulsion (AFM) between tip and surface

0.5 nm

0.1 nm

SEM: scanning electron microscopy

0.5 nm

4 nm

0.3 m

2 m

Secondary electron emission caused by focused electron beam is measured and spatially imaged

SPR: surface plasmon Measurement change in the resonance refractive index in evanescent field generated surface

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material surface under such conditions must be made very carefully because the actual surface has hydrophilic and hydrophobic components. Surfaces readily contaminate. Under ultra-high vacuum conditions, this contamination can be retarded. However, in view of the atmospheric pressure conditions under which all biomedical devices are used, we must learn to live with some degree of contamination. The significant questions here are whether we can make devices that have a constant, controlled level of contamination and can avoid undesirable contamination. This is critical in order that a laboratory experiment on a biomaterial generates the same results when repeated after one day, one week, or one year, and that a biomedical device performs for the physician in a constant manner over a reasonable operating life. Some analytical methods also apply to the characterization of a biological surface that forms an interface with an artificial material. Compared with an artificial surface, a biological surface is more complex, flexible, and easier to contaminate. To characterize the biological surface, we must then be cognizant of the circumstances under which the biointerface is used. A wide variety of methods has therefore to be applied for understanding the reliable surface properties of materials. Here, a few typical methods for the surface analysis of biomaterials are introduced.

14.3.2 Contact angle method The force balance between the liquid-vapor surface tension ( LV ) of a drop of liquid and the interfacial tension between a solid and the drop ( SL ), which is manifested through the contact angle () of the drop with the surface (Fig. 14.2), can be used to characterize the energy of the surface ( SV ). The basic relationship describing this force balance is:

SV ˆ SL ‡ LV
cos 

14:1

The energy of the surface, which is directly related to its wettability, is a useful parameter that has often correlated strongly with biological interaction. Unfortunately, SV cannot be obtained directly since this equation contains two unknowns, SL and SV . Therefore, SV is usually approximated by the

14.2 Water contact angle measurement using a sessile drop technique.

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14.3 Schematic representation of corona irradiation apparatus.

Zisman method for obtaining the critical surface tension.7 Experimentally, there are several ways to measure the contact angle. From this measurement, a fundamental characterization of material surface can be obtained. The corona discharge treatment is one of the most robust methods for surface modification of polymer materials to control surface wettability and produce graft polymers. Lee and co-workers prepared a wettability gradient surface on conventional polymer materials and investigated protein or cell/surface interactions.8 Figure 14.3 is a schematic diagram showing the corona discharge apparatus for the preparation of a gradient polyethylene sheet. Peroxide is produced by the corona discharge treatment, and this peroxide transfers to various functional groups with oxygen. Figure 14.4 shows the water contact angles of the PE surface after treatment with corona irradiation by changing the energy along the sample distance. The water contact angle is decreased with an increase in the energy of the corona irradiation.

14.4 Change in water contact angle of PE surface by corona treatment.

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14.3.3 X-ray photoelectron spectroscopy (XPS)9 X-ray photoelectron spectroscopy (XPS) is generally regarded as an important and key technique for the surface characterization and analysis of biomedical polymers. This technique, also called ESCA, provides a total elemental analysis, Ê of any solid surface that is except for hydrogen and helium, of the top 10±200 A vacuum stable or can be made vacuum stable by cooling. Chemical bonding information is also provided. Of all the presently available instrumental techniques for surface analysis, XPS is generally regarded as being the most quantitative, the most readily interpretable, and the most informative with regard to chemical information. For these reasons, it has been highly recommended and used by biomedical researchers for the analysis of medical materials. The basic principle of XPS is the photoelectric effect, the phenomenon for which Einstein received the Nobel Prize. Three possibilities are evident. 1. 2. 3.

The photon may traverse through the atom without significantly interacting with either the orbital electrons or the nucleus. The photon may be scattered by the atomic orbital electron resulting in a partial loss of photon energy. This process is called Compton scattering. The photon may interact with the atomic orbital electron such that there is total and complete transfer of the photon's energy to the electron. This is the basic process in XPS.

Given that the photon energy is greater than the binding energy of the electron in the atom, the electron is then ejected from the atom with a kinetic energy approximately equal to the difference between the photon energy and the binding energy (Fig. 14.5). Therefore, the basic equation for XPS is: Eb ˆ h ÿ Ek

14:2

where Eb is the electron binding energy, Ek is the electron kinetic energy measured by the instrument, and h is the photon energy (h is Planck's constant and  is the X-ray frequency). All energies are usually expressed in electron volts (eV). Measuring the kinetic energy enables calculation of the binding energy. By knowing the binding energy, we can identify the atom.

14.5 Schematic views of XPS analysis and interaction of X-ray photon of energy h with an atomic orbital electron.

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14.6 XPS spectra of the gradient polymer surfaces at positions 0.5 cm, 2.5 cm, and 4.5 cm from the starting point of corona treatment.

The peroxides produced on a polymer substrate by corona irradiation (Fig. 14.3) work as an initiator for free-radical polymerization. When the coronairradiated polymer film is soaked in a monomer solution, the graft polymers were obtained from the film surface. Figure 14.6 shows the XPS spectra of C1s and P2p on a gradient poly(2-methacryloyloxyethyl phosphorylcholine) (PMPC) grafted polyethylene surface.10 The concentration of the MPC unit on the surface increased with an increase in the distance from the starting point of the corona irradiation. That is, a phosphorus peak attributed to the phosphorylcholine group in the MPC unit was found at 134.5. Moreover, a carbonyl peak was seen at 288.5 eV because it was clear compared with the untreated PE surface. The protein adsorption and cell adhesion decreased with an increase in the density of the grafted poly(MPC).

14.3.4 Scanning probe microscopy (SPM)11 A number of methods that provide information about the structure of a solid surface, its composition, and current oxidation states have come into use. The recent upsurge of activity in scanning probe microscopy has resulted in

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14.7 Schematic diagram of the scanning tunneling microscope (STM) and atomic force microscope (AFM).

investigation of a wide variety of surface structures under a range of conditions. The ability to control the position of a fine tip in order to scan surfaces with subatomic resolution has brought scanning probe microscopy to the forefront in surface imaging techniques. There are two primary techniques; scanning tunnel microscopy (STM) and atomic force microscopy (AFM), as illustrated in Fig. 14.7. STM relies on measurement of the experimentally decaying tunneling current between a metal tip and the conducting substrate. Since its development in the early 1980s and the recognition of its inventors with the presentation of the 1986 Nobel Prize, STM has found wide use in studies of both inorganic and organic materials. Supporting a fine tip, a piezoelectrically positioned cantilever spring provides the means of measuring surface forces in the range of 10ÿ13 to 10ÿ6 N. The AFM measures deflections in the cantilever due to capillary, electrostatic, van der Waals, and frictional forces between the tip and the surface. Not limited to conductive surfaces, AFM measurements can also be made on organic and inorganic surfaces and on surfaces immersed in liquids. Conical tips of silicon have points of 5 to 50 nm (radius of curvature). However, numerous probes have been used including attaching a colloidal particle of several micrometers to a cantilever, as described by Butt and co-worker12 and Ducker and co-worker.13 Since AFM measures force, it can be used with both conductive and nonconductive specimens. Since force must be applied to bend a lever, AFM is subject to artifacts caused by damage to fragile structures on the surface. Both STM and AFM methods can function well for specimens under water, in air, or under a vacuum. For exploring biomolecules or mobile organic surfaces, the `pushing around' of structures by the tip is a significant concern. With both methods, it is difficult to achieve good-quality, reproducible images of organic substrates. However, some of the successes to date are exciting enough that the future of these methods in biomedical research is ensured. Figure 14.8 is an AFM image of a DPPC liposome adsorbed on a 2methacryloyloxyethyl phosphorylcholine (MPC) polymer surface. Using AFM, we can observe biomolecules and their association with the native structure.14

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14.8 AFM image of liposomes adsorbed on the MPC polymer.

14.4 Design for non-biofouling surface An assortment of materials have been used for manufacturing medical devices including artificial organs that are exposed to blood. However, the only polymers presently used for this type of application are conventional materials such as poly(vinyl chloride) (PVC), polyethylene(PE), poly[methyl methacrylate (MMA)], segmented polyether urethane (SPU), poly(dimethylsiloxane), poly(tetrafluoroethylene), cellulose, and polysulfone (PSF). Because nonspecific protein adsorption occurs on these materials, they cannot reduce the level of reactions to foreign bodies. Therefore, drug infusion is required during clinical treatments using these medical devices to prevent reactions to foreign bodies. To reduce the level of bioreaction on a surface, polymeric surfaces, which can reduce nonspecific protein adsorption, that is, `nonfouling surfaces' have been studied. The molecular design of nonfouling polymers for biomedical

Poly(2-hydroxyethyl methacrylate) Poly(acrylamide) Block-type copolymer Graft-type copolymer Segmented polyurethane Poly(ethylene glycol)

Hydrophilic surface zero interfacial free energy

Heterogenic surface microdomain concept

Molecular cilia mechanism

Negative charged surface Sulfonated polymer - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - Synthetic polymer Pseudomembrane formation Expanded PTFE + Poly(ethylene terephthalate) Biologically active Immobilization of bioactive protein polymer Heparinization Heparine releasing polymer Heparinized polymer Urokinase Thrombomoduline - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - Synthetic polymer Albumin adsorbed surface Polyurethane with alkyl group + Biological Phospholipid adsorbed surface Phospholipid polymer molecules biomembrane-like surface formation

Polytetrafluoroethylene(PTFE) Polydimethylsiloxane Polyethylene

Hydrophobic surface zero critical surface theory

Synthetic polymer

Typical polymer

Basic concept

Type of materials

Table 14.4 Surface immobilizations for blood contacting materials

s

Biological approach

Biochemical approach

Interfacial chemical approach

Physicochemical approach

s s

s

Approach

ss

s

s

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14.9 Chemical structure of MPC.

applications is classified into the four categories listed in Table 14.4. One of the most effective methods for producing a nonfouling surface is the modification of conventional materials with polymers having a phospholipid polar group mimicking a biomembrane surface. In 1978, Nakabayashi designed a methacrylate monomer with a phospholipid polar group, 2-methacryloyloxyethyl phosphorylcholine (MPC) to obtain new medical polymer materials (Fig. 14.9).15 However, at that time, the purity and yield of the MPC was not sufficient to evaluate their functions. Ishihara improved the synthetic route of the MPC16 and prepared a wide variety of polymers containing the MPC unit.17±19 The homopolymer of the MPC is soluble in water and the solubility of the MPC polymers can easily be controlled by changing the structure and fraction of the comonomer. Many papers have described the effect of the MPC moiety on the nonfouling property using poly(MPC-co-n-butyl methacrylate(BMA)) (PMB). Figure 14.10 shows experimental columns containing polymer-coated poly(methyl methacrylate) beads and micrographs of the bead surface after contact with human blood without an anticoagulant. On the hydrophobic poly[nbutyl methacrylate (BMA)], many adherent cells can be observed. Moreover, the activation and aggregation of the cells and clot formation can be observed on the

14.10 Blood compatibility test on PBMA and PMB30 with microsphere column method.

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14.11 Amount of protein adsorbed on polymer surfaces from human plasma.

poly(BMA). In contrast, poly(MPC-co-BMA) (PMB) with a 30 mol% MPC unit (PMB30) effectively suppressed cell adhesion. The blood compatibility of MPC polymer is due to the surface property that reduces nonspecific protein adsorption. The protein adsorption-resistance on MPC polymer surface has been studied with consideration of not only the amount of adsorbed proteins but also of the species of the proteins. Figure 14.11 shows the amount of protein adsorbed on the PMB, poly(HEMA), and poly(BMA) after contact with plasma for 60 min.20 On the poly(BMA), many more proteins were adsorbed compared with those on the poly(HEMA) and the PMB. The amount of proteins adsorbed on the PMB decreased with an increase in the amount of MPC in the polymer. The species and distribution of the proteins adsorbed on the PMB were also determined by gold colloid- and radio-labeled immunoassays. 21 These experiments demonstrated that the PMB could reduce plasma protein adsorption nonspecifically. The thrombus formation on the conventional polymeric materials occurred through the multilayers of plasma proteins denaturated by contact with the surfaces. The secondary structure of bovine serum albumin (BSA) and bovine plasma fibrinogen (BPF) adsorbed on the PMB was evaluated by circular dichroism (CD) spectroscopy.22,23 Figure 14.12 shows the CD spectra of BSA in PBS and that adsorbed on the polymer surface. For the BSA in PBS, the mean molecular residual ellipticity, [], was a large negative value at 222 nm. The CD spectrum of BSA adsorbed on

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14.12 CD spectra of BSA in PBS and that on adsorbed polymer surfaces. (A): poly(HEMA), (B): PMB10, (C): PMB30, (D): BSA/PBS.

PMB was almost the same as that in PBS. The negative ellipticity at 222 nm of BSA adsorbed on the MPC polymers increased with a decrease in the ratio of MPC, then neared zero in the case of BSA adsorbed on poly(HEMA). We were able to find the same tendency for BPF. Calculation of the -helix content of BSA and BPF revealed that the PMB could effectively suppress the conformational change of the proteins even when the proteins were adsorbed on the surface. Conversely, it could be seen that the -helix content of both proteins adsorbed on poly(HEMA) decreased significantly. It is thought that the MPC polymer suppresses protein adsorption. Recently, it was reported that the structure of the water absorbed in/on the polymer materials influences protein adsorption on their surfaces. The water structure in hydrated polymers was determined to enhance the resistant properties of protein adsorption on surfaces containing phospholipids. Park et al. proposed a very important model for protein adsorption on polymer surfaces. The adsorption of proteins on a polymer surface via hydrophobic interaction requires an exchange of bound water between the protein and the surface.24 Therefore, the amount of bound water might be the important parameter in understanding protein adsorption. Tsuruta reported that the random networks of water molecules on material surfaces were very important in explaining the protein adsorption that occurs on them.25 Protein adsorption processes are considered to start with protein trapping by random networks of water molecules on the material surface. The material surface, which cannot construct hydrogen bonding with water, will then reduce protein adsorption. Table 14.5 lists the free water concentration in a hydrated

404

Surfaces and interfaces for biomaterials Table 14.5 Characteristics of hydration state of polymer Poly(HEMA) Heqa Free water fraction at Heq at H=0.36 a

Heq ˆ

0.40 0.34 0.28

PMB 10

30

0.23 0.25

0.84 0.84 0.69

Weight of water in the polymer membrane at 25 ëC Weight of polymer membrane saturated with water

polymer membrane with a 0.36 water fraction determined by differential scanning calorimetry.23 The fraction of free water (not bound water) in the MPC polymer was 0.65, which was found to be significantly higher than that in poly(HEMA), which was 0.28. Furthermore, the structure and hydrogen bonding of water near the PMB were analyzed in their aqueous solutions and in thin films with contours of O-H stretching of Raman and attenuated total reflection infrared (ATR-IR) spectra, respectively.26,27 The relative intensity of the collective band (C value) corresponding to a long-range coupling of O-H stretchings of the Raman spectra for the aqueous solution of PMB was very close to that for pure water, which is in contrast with the smaller C value in the aqueous solution of ordinary polyelectrolytes. A similar tendency was also observed on hydrated thin polymer films. These results suggest that PMB does not significantly disturb hydrogen bonding between water molecules in either the aqueous solution or the thin film systems. The equilibrium amount of proteins, bovine serum albumin (BSA), and bovine plasma fibrinogen (BPF) adsorbed on the polymer surface was measured and represented with a free water fraction in the hydrated polymers, as shown in Fig. 14.13.28 The amount of both proteins adsorbed on poly(HEMA), poly(acryl amide (AAm)-co-BMA), and poly(N-vinylpyrrolidone(VPy)-co-BMA) were larger than those on the poly(MPC-co-dodecyl methacrylate)(PMD) and PMB. It was reported that the theoretical amount of BSA and BPF adsorbed on the surface in a monolayer state are 0.9 and 1.7 g/cm2, respectively. On the surfaces of MPC polymers, the amount of adsorbed proteins was less than these theoretical values. A schematic representation of the nonfouling property of PMB surfaces is shown in Fig. 14.14. When the PMB comes in contact with biofluids containing proteins, the hydrated surface of the PMB has a high free-water content and cannot be recognized by proteins. Cell adhesion is also reduced on these surfaces because there is no adhesive protein on the surfaces. Recently, it was found that the MPC polymer surface could reduce not only biofouling but also the inflammatory responses of cells in contact with the surface. Therefore, compared with other nonfouling surfaces, the MPC polymer surface might offer many advantages for biomedical materials.

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14.13 Relationship between free water fraction in hydrated polymer membrane and amount of proteins adsorbed on the polymer. :BSA, [BSA] in PBS ˆ 0.45 g/dl, :BPF, [BPF] ˆ 0.30 g/dl.

14.14 Possible mechanism of nonfouling property on MPC polymer surface.

MPC is a robust compound for surface modification to improve biocompatibility because it can be used in a wide variety of technologies. Figure 14.15 shows the surface modification of conventional polymer materials with MPC. Blood-compatible biomedical devices employing these technologies have been explored.

14.5 How to connect tissues with biomaterials There is a great demand for connecting tissues to devices, for artificial organs, and for biomaterials to support organ functions and injured and/or unwell tissues. It is not so difficult to unite natural tissues, which could heal by

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Surfaces and interfaces for biomaterials

14.15 Various MPC polymer immobilizations on conventional polymer surface.

themselves. However, this is not always possible and we have to assist organ and tissue functions with biomaterials. The ability to introduce blood to a dialyzer three times a week is required for patients with chronic kidney ailments to keep them alive. Artificial blood vessels must be sutured to natural blood vessels because there is no suitable adhesive to connect biomaterials and tissues. Infection at the interface is a severe problem in medical treatment. Living tissues generally reject biomaterials in an attempt to avoid adverse effects. Nevertheless, we are attempting to apply biomaterials to remedy an organ failure such as the use of an artificial kidney. Adequate blood access is a critical problem for patients with chronic kidney disease. In several senses, biocompatibility is required at the tissue/biomaterial interface. When we consider the mechanism of pannus formation at the natural vessel side of connected artificial and natural blood vessels working for a long time, the differences in their biomechanical properties must be considered. The pulsing of blood vessels causes severe stress when the mechanical properties are different. The natural vessels change to minimize the adverse effect of the artificial vessels and their thickness increases to accommodate the situation. `Adhesive' is a general term that covers designations such as cement, glue, paste, fixative, and bonding agents used in the many areas of adhesive technology. Adhesive systems consist of one- or two-part organic and/or inorganic formulations that set or harden through the action of several different mechanisms. The applications of adhesive biomaterials range from soft tissue adhesives used both externally to temporarily attach accessory devices such as colostomy bags and internally for wound closure and sealing, to hard tissue adhesives used to bond prosthetic materials to teeth and bone on a more permanent basis. All of these physiological environments are hostile, and a major problem in the formulation of medical and dental adhesives is to develop

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a material that will be easy to manipulate, interact intimately with the tissue to form a suitable bond, and be biocompatible. Over the past two decades, more success at a clinical level has been achieved in bonding hard tissues than soft tissues. The tooth surface is an important surface, which should be given special note. Dental biomaterials are applied to rehabilitate decayed areas formed on dental hard tissues, because these tissues do not heal and regenerate. Completion of wound healing is essential to inhibit infection at the surface of the injured tissue before applying any biomaterials. Unfortunately, teeth do not heal by themselves because there are no blood vessels to initiate the wound-healing process. Dentists have been affixing restorations to prepared cavities and abutments as ordinary dental treatment for a long time. However, the treatment often does not survive on the tissues and the prostheses soon detach. It has been hypothesized that a good adhesive for dental hard tissues could provide good resolution in improving dental treatment by increasing the retentive force combined with the adhesive force. However, connecting natural tissues, even dental hard tissues, to biomaterials is not so easy. Dentists believe that the detachment of a restoration occurs by `microleakage' and `secondary caries' but they could not explain these mechanisms exactly. They misunderstood that dental cements cured by acid and base reactions are soluble in saliva. Therefore, they tried to minimize the dissolution of cured zinc phosphate cement luted on the hard tissues and accurately fabricated restorations with the prepared tissues by decreasing the thickness of the cement. They developed zinc carboxylate cements and glass-ionomer cements to decrease the solubility in saliva. They should understand that the cements dissolved if the abutments were dentin but not if the abutments were enamel. There are big differences between the two types of tissue. Cured cements can intimately attach to enamel etched with acidic cement pastes before they cure. In contrast, they are attached to dentin demineralized with acidic cement pastes before they have cured. The demineralized dentin is permeable to saliva, which dissolves the cement. In 1955, Buonocore developed a method of bonding self-cured acrylic resin to etched enamel. This technology was introduced in orthodontic treatment to affix plastic and metal brackets directly to provide stress on teeth in order to align them. Bonding to enamel is believed to be not as difficult as bonding to dentin, and the etching of the enamel gives a good bonding surface. Enamel does not contain as much water and peptides as dentin does; the principal constituents are hydroxyapatite crystals. Tag formation on the etched enamel surface gives good mechanical retention, which provides bond strength to the substrate. Therefore, stronger etching agents are used to make bigger tags to create stronger bonds to enamel. However, aggressive etchants weaken enamel. The bonding mechanism is explained in the illustration in Fig. 14.16.29 During studies to improve the direct bonding system in orthodontic treatment, it

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14.16 Schematic diagram of hybrid concept for resin bonding to enamel.

was found that the length of the tags changes with the adhesive used. Adhesives containing a hydrophobic and hydrophilic methacrylate appeared to penetrate beyond the etched surface and encapsulate the prisms. The hybrid zone does not increase the bonding strength significantly, but does promote stability due to a supplemental thin layer of resin-reinforced tissue. This helps minimize the adverse effect of etching the enamel and gives the enamel acid resistance. Several mechanisms have been proposed for bonding to dentin. Fixation of restorative materials had been performed by curing fluid substances on the surface of the tooth, which is on the outside of the tooth. Many dental researchers thought adhesive bonding was the result of surface phenomena. Consequently, wettability and chemical reaction were thought to be important in creating bonds to dentin. Unfortunately, these considerations have not worked in developing reliable systems for bonding to dentin. In 1982, Nakabayashi proposed a new mechanism to enable the bonding of resin to dentin;30 this system was widely accepted in the 1990s. The system is described as being micromechanical at the molecular level. Monomers having both hydrophobic and hydrophilic groups are impregnated by the exposed collagen of demineralized superficial dentin, which has a good capacity to accept the diffusion of monomers. Their polymerization in situ forms copolymers entangled with the collagen fibers and encapsulated hydroxyapatite crystalline. The entanglement locks the dentin with cured resin, which affords

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the bonding strength. It is the generation of a `hybridized dentin,' which is a transitional zone of resin-reinforced dentin, sandwiched between cured resin and the unaltered dentin substrate, and a mixture of dentinal components and cured resin at the molecular level. It is a kind of functional graded material. The hybrid layer is resin-infiltrated enamel, dentin, or cementum. The chemical and physical properties of these zones are quite different from those of the original tooth structure because the tissue has been partially demineralized and then impregnated with resin. The resulting structure is neither resin nor tooth but a hybrid of the two. It is not located on the surface but prepared within the subsurface of the substrates. The zones are so-called tissue-engineered demineralized dentin with polymers. The mineral phase of the hard tissue is purposely dissolved to create a diffusion pathway to the monomers. This matrix is then infiltrated with monomers to intentionally change their physical and chemical properties. The clinical success of pit and fissure sealants on enamel indicates how acid resistant these resin-enamel hybridizations can be over many years. The infiltrating resin actually envelops apatitic crystallites in enamel to improve their acid resistance. This is made possible by the pretreatment of enamel with acids, which increases the surface roughness on a microscopic scale. Indeed, resinbonded enamel contains micro-tags of resin prepared between the crystalline and enamel rods etched at a diameter as small as the molecular level, as illustrated in Fig. 14.16. It is essential to know the importance of hybridized dentin in dental treatment more than the bonding of restorations to tooth structures. We have believed that caries are formed by infections of acid-producing microorganisms living in our mouths. This is true in a sense, but we must note that these microorganisms do not infect dental tissues from the beginning of tooth development. Acids, mostly lactic acid and their related products, react with basic calcium phosphate, hydoxyapatite, and dissolve the tissues. It is a simple neutralization and demineralization of enamel and dentin. The demineralized enamel surface, without visible caries, is restored by the recrystallization of the hydroxyapatite present in the saliva. If the rate of demineralization does not exceed that of recrystallization, the enamel is sufficiently stable dynamically. However, if the rate of the former is faster than that of the latter, caries should develop. Consequently, it is very important to clean our teeth after eating food, which helps microorganisms form a plaque biofilm. The development of caries is evidence of a good acid delivery system to the enamel. Enamel is a very important tissue that protects the health of our teeth by acting as a barrier to protect the dentin and underlying pulp. It does this by inhibiting the invasion of lactic acid to the dentin to form dentin caries. Once dentin is exposed, it is impossible for dentists to heal it using dental materials, which has been a basic misunderstanding. The pulp in exposed dentin must accept the invasion of stimuli and microorganisms, which irritate the pulp. To

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prevent septicemia, sometimes we have to kill the tooth if the pulp must be extracted. Non-vital teeth constitute dead tissue. We have to prevent it by applying artificial enamel at the early stages of caries development and inhibit the invasion of lactic acid as our teeth cannot inhibit it. The healing of wounds of exposed dentin caused by caries is impossible and we must naturally extract the tooth. Although, dentists have tried to heal caries, this is impossible, as mentioned before. Blood coagulation after bleeding could initiate gingival wound healing and protect us from further infection. As is enamel, hybridized dentin is an impermeable barrier to lactic acid and can inhibit further demineralization of exposed dentin. We have been suffering secondary caries and detachment of prostheses affixed to abutments. Dental researchers have worked hard to inhibit microleakage by several methods but resolution was difficult. The meaning of microleakage, which damages restored teeth, is the marginal leakage of lactic acid, the principal product of acidproducing microorganisms in the mouth, and demineralization of abutments by the acid. We could say that hybridized dentin is artificial enamel that can protect dentin from further demineralization caused by acid. The preparation of hybridized dentin could be a man-made process of wound healing, pseudowound healing, of exposed dentin because it can protect further demineralization leaving only healed tissue. We could then easily provide rehabilitation by affixing prostheses to reconstruct the esthetics and functions of lost tissues. Dentists could attach the restoration to the top of the hybridized dentin because the top surface has good affinity for attachment with resin composites and resin cements. The interaction of tissues and biomaterials at biologic interfaces is extremely important but it is very difficult to connect natural tissues with artificial materials. The bonding of resins to mineralized tissues provides a good example of such an interface. The three major hard tissues are enamel, dentin, and bone. Unlike bone, the former two, the hard tissues of the teeth, do not regenerate. When biocompatible materials, such as dental implants, are placed close to bone, hydroxyapatite prepared by the surrounding bone fills the intervening space and a connection is formed. In contrast, when such a material is placed close to a tooth, no such connection develops because there is no biologically active soft tissue between the two structures to promote regeneration (Fig. 14.17). Unfortunately, hybridization of bone with impregnated adhesives killed bone-inducing cells by their encapsulation, and the bonding mechanism could not be applied to develop an adhesive bone cement between bone and prostheses.

14.6 Conclusion Surface design and characterization as well as mechanical properties are quite important in the preparation of biomaterials. The moment a living body is

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14.17 Interactions of biocompatible materials with tooth, bone, and soft tissue.

exposed to a material surface, it determines whether the material can be accepted. Requirements of biomaterial surfaces vary depending on their use and the environment in which they are used. However, biocompatibility is a necessary element for every biomaterial. Concerning biological structure and function, and applying various analyses for surface characterization may contribute to the exploration of suitable material surfaces.

14.7 References 1. Ratner B D (1996), 'Surface properties of materials,' in Ratner B D, Hoffman A S, Schoen F J, and Lemons J E, Biomaterials Science, San Diego, Academic Press, 21± 34. 2. Schoen F J and Anderson J M (1996), `Host reactions to biomaterials and their evaluation,' in Ratner B D, Hoffman A S, Schoen F J, and Lemons J E, Biomaterials Science, San Diego, Academic Press, 165±173. 3. Williams D F (1987), `Definitions in Biomaterials'. Proceedings of a Consensus of the European Society for Biomaterials, Chester, England, March 3±5 1986, New York, Elsevier. 4. Lee J W and Gardella J A Jr. (2002), `Surface perspectives in the biomedical applications of poly(alpha-hydroxy acid)s and their associated copolymers,' Anal Bioanal Chem, 373, 526±573. 5. Brash J L and Horbett T A (1987), Proteins at Interfaces ± Physicochemical and Biochemical Studies, ACS Symposium Series 343, Washington, DC, American Chemical Society. 6. Brash J L and Horbett T A (1995), Proteins at Interfaces II ± Fundamentals and Applications, ACS Symposium Series 602, Washington, DC, American Chemical Society. 7. Zisman W A (1964), `Relation of the equilibrium contact angle to liquid and solid constitution,' in Fowkes F M, Contact angle, wettability and adhesion, ACS Advances in Chemistry Series 43, Washington, DC, American Chemical Society, 1±51.

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8. Lee J H and Lee H B (1993), `A wettability gradient as a tool to study protein adsorption and cell adhesion on polymer surfaces,' J Biomater Sci Polym Ed, 4, 467±81. 9. Andrade J D (1985), Surface and Interfacial Aspects of Biomedical Polymers, vol. 1, Surface chemistry and Physics, New York, Plenum Press. 10. Iwasaki Y, Ishihara K, Nakabayashi N, Khang G, Jeon J H, Lee J W, and Lee H B (1998), `Platelet adhesion on the gradient surfaces grafted with phospholipid polymer,' J Biomater Sci Polym Ed, 9, 801±816. 11. Adamson A W and Gast A P (1997), Physical chemistry of surfaces, New York, John Wiley & Sons, Inc. 12. Butt H J, Jaschke M, and Ducker W (1995), `Measuring surface forces in aqueous electrolyte solution with the atomic force microscope,' Bioelectrochem Bioenerg, 38, 191±201. 13. Ducker W A, Xu Z (1994), Israelachvili J N, `Measurement of Hydrophobic and DLVO Forces in Bubble-Surface Interactions in Aqueous Solutions,' Langmuir, 10, 3279±3289. 14. Iwasaki Y, Tanaka S, Hara M, Ishihara K, and Nakabayashi N (1997), `Stabilization of liposome attached on polymer surface having phosphorylcholine group,' J Colloid Interface Sci, 192, 432±439. 15. Kadoma Y, Nakabayashi N, Masuhara E, and Yamauchi J (1978), `Synthesis and hemolysis test of the polymer containing phosphorylcholine groups,' Koubunshi Ronbunshu (Jpn J Polym Sci Tech), 35, 423±427. 16. Ishihara K, Ueda T, and Nakabayashi N (1990), `Preparation of phospholipid polymers and their properties as hydrogel membrane,' Polym J, 22, 355±360. 17. Ueda T, Oshida H, Kurita K, Ishihara K, and Nakabayashi N (1992), `Preparation of 2-methacryloyloxyethyl phosphorylcholine copolymers with alkyl methacrylates and their blood compatibility,' Polym J, 24, 1259±1369 (1992). 18. Ishihara K, Tsuji T, Kurosaki K, and Nakabayashi N (1994), `Hemocompatibility on graft copolymers composed of poly(2-methacryloyloxyethyl phosphorylcholine) side chain and poly(n-butyl methacrylate) backbone,' J Biomed Mater Res, 28, 225± 232. 19. Ishihara K, Inoue H, Kurita K, and Nakabayashi N (1994), `Selective adhesion of platelets on a polyion complex composed of phospholipid polymers containing sulfonate groups and quaternary ammonium groups,' J Biomed Mater Res, 28, 1347±1355. 20. Ishihara K, Oshida H, Endo Y, Ueda T, Watanabe A and Nakabayashi N (1992), `Hemocompatibility of human whole blood on polymers with a phospholipid polar group and its mechanism,' J Biomed Mater Res, 26, 1543±1552. 21. Ishihara K, Ziats N P, Tierney B P, Nakabayashi N, and Anderson J M (1991), `Protein adsorption from human plasma is reduced on phospholipid polymer,' J Biomed Mater Res, 25, 1397±1407. 22. Ishihara K, Ueda T, Saito N, Kurita K, and Nakabayashi N (1991), `Suppression of protein adsorption and denaturation on polymer surface with phospholipid polar group,' Seitai Zairyou (J J Soc Biomat), 9, 25±31. 23. Ishihara K, Nomura H, Mihara T, Kurita K, Iwasaki Y, and Nakabayashi N (1997), `Why do phospholipid polymers reduce protein adsorption?' J Biomed Mater Res, 39, 323±330. 24. Lu D R, Lee S J and Park K (1991), `Calculation of solvation interaction energies for

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protein adsorption on polymer surfaces,' J Biomater Sci Polymer Ed, 3, 127±147. 25. Tsuruta T (1996), `Contemporary topics in polymeric materials for biomedical applications,' Adv Polym Sci, 126, 1±51. 26. Kitano H, Sudo K, Ichikawa K, Ide M, Ishihara K (2000), `Raman spectroscopic study on the structure of water in aqueous polyelectrolyte solutions,' J Phys Chem B, 104, 11425±11429. 27. Kitano H, Imai M, Mori T, Gemmei-Ide M, Yokoyama Y, Ishihara K (2003), `Structure of water in the vicinity of phospholipid analogue copolymers as studied by vibrational spectroscopy,' Langmuir 19, 10260±10266. 28. Ishihara K (2000), `Bioinspired phospholipid polymer biomaterials for making high performance artificial organs,' Sci Tech Adv Mater, 1, 131±138. 29. Nakabayashi N and Pashley D H (1998), Hybridization of dental Hard Tissues, Tokyo, Quintessence publishing Co., Ltd. 30. Nakabayashi N, Kojima K, Masuhara E (1982), `The promotion of adhesion by the infiltration of monomers into tooth substrates,' J Biomed Mater Res, 16, 265±73.

15 Blood flow dynamics and surface interactions W V A N O E V E R E N , University of Groningen, The Netherlands

15.1 Clinical application and problems of medical devices in contact with blood Application of biomaterials in direct blood contact results in activation of the blood coagulation system and in an inflammatory reaction. These responses of blood are due to the natural response of the host defence mechanism against foreign surfaces. Inadequate control by natural inhibitors results in pathological processes, such as microthrombi generation or thrombosis, bleeding complications, haemodynamic instability, fever, edema, and organ damage. These adverse events become manifest during prolonged and intensive foreign material contact with vascular implants and extracorporeal blood circulation.1-7 Medical Device Alert in the UK has shown that the number of adverse events induced by CE-marked products is increasing (Fig. 15.1). More than one quarter of those adverse events is due to blood contacting devices. FDA reports indicate a doubled or even tripled increase of thrombotic incidences of implants from 1998 to 2003 (Fig. 15.2). It is hard to speculate on a direct relation with haemocompatibility, but consensus exists about the important role of poor haemocompatibility in direct and sustained adverse reactions. Some of the most studied effects of biomaterials in contact with blood will be discussed in this section.

15.1.1 Small implants in the blood vascular system: stents and vascular grafts After balloon dilatation of narrowed small diameter arteries, stents are frequently applied to maintain the lumen open. The basis material of stents is a metal, such as stainless steel, tantalum, nitinol, based on the mechanical properties of these metals to support the vascular wall with minimal occlusion of side branch capillaries. The occurrence of in-stent subacute thrombosis has been reduced, but still remain as a worrying complication because of its strong impact on short-term mortality.8 Due to the refinement of adjunctive antiplatelet

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15.1 Reports of incidents with medical devices in the UK indicate a significant increase of CE marked products, partly due to an increased use of such CE products, but also indicating the admission to the market of malfunctioning medical devices. Adverse events in circulating blood often cause serious health problems, which may explain the high frequency (28%) of blood contacting devices in this figure (adapted from presentation by Tony Sant, Manager Medical Devices Agency, MDA liaison officer conference 6 March 2002).

treatment9 and the establishment of the most appropriate ways for stent utilisation,10 reduced thrombosis was observed around the millennium change. However, an increase of thrombotic incidents has been observed since, due to increased stent utilisation, and also related to new types of bioactive coatings, aiming for reduction of late reappearance of a coronary stenosis at the site of intervention (restenosis). In-stent subacute thrombosis and restenosis arise partly from the same source. The immediate local deposition of platelets and leukocytes onto the bare material surface can lead to abrupt lumen obliteration and enhances the release of cytokines and growth factors, thus promoting intimal hyperplasia of smooth muscle cells into the vessel wall.11 Intimal

15.2 Thrombotic events are one of the most obvious side effects of implants in the vascular system. According to the FDA (USA) the number of thrombotic events reduced in three types of devices in the years 1996 to 2000. This can be explained by the use of aggressive antithrombotic drugs (on the cost of increased bleeding complications). However, new types of catheters and stents have induced a significant increase of thrombosis in recent years (adapted from: MAUDE database search, Center for devices and radiological health, FDA).

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hyperplasia is an important reason for late stent or graft failure. It is recognised as intrinsic obstructive lesion after many forms of arterial wall injury, for instance, in coronary and superficial femoral artery angioplasty, endarterectomy, arteriovenous fistulae for haemodialyis and homograft transplanted organs.12 Intimal hyperplasia occurs physiologically in closure of the ductus arteriosus after birth.13 Implantation of a medical device into the arterial circulation leads to endothelial denudation, which is immediately followed by deposition of platelets and leucocytes. Besides damage to the endothelium, high arterial pressure and flow cause damage to the medial layer of the blood vessel. Approximately six hours after implantation leucocytes infiltrate the vessel media. This medial layer is to a certain extent damaged resulting in the death of smooth muscle cells. Both dead and injured endothelium and medial smooth muscle cells are able to release growth factors. One of these released growth factors is basic fibroblast growth factor (bFGF) which stimulates the proliferation of endothelial cells and of smooth muscle cells.14 Besides proliferation, the synthetic smooth muscle cells produce extracellular matrix resulting in medial thickening.15 Another important growth factor is platelet derived growth factor.16 Particularly biomaterial implants with poor haemocompatibility and poor rheology will cause platelet activation and subsequent release of platelet derived growth factor. The triggers for the formation of intimal hyperplasia that have been further defined are injury, circulating blood components, and haemodynamics.17±19

15.1.2 Pharmacologic treatment for heart valves, coronary stents and vascular prostheses The most common platelet aggregation inhibitor is aspirin. Aspirin and other non-steroid anti-inflammatory drugs irreversibly acetylate a serine residue in the active site of cyclo oxygenase blocking the formation of thromboxane. Aspirin has no capacity to block the release of platelet derived growth factor nor the capacity to block the first wave of ADP induced platelet aggregation,20 so the effect of platelet aggregation inhibitors is theoretically low. In experimental models of vein grafting, conflicting results are present on the reduction of intimal hyperplasia using platelet aggregation inhibitors.21,22 The currently recommended antiplatelet treatment after stenting (Aspirin + Ticlopidine or Clopidogrel) has been standardised after experimental and clinical observations with the first generation of stents, and since then universally applied. However, it has to be proven that this combination of antiplatelet drugs is suitable for any procedure. Besides, it seems likely that, in particular situations (optimal stent expansion, less device-induced thrombogenicity), patients can profit by a sole, though adequate, antiplatelet treatment.23 Furthermore, the population of cardiopathic patients submitted to percutaneous interventions

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evolves continuously and dramatically fast. Lately stenting procedures have been safely and effectively implemented in the treatment of patients suffering from acute coronary syndromes (unstable angina, acute myocardial infarction). Many pivotal trials have demonstrated that, in such patients, the blockade of activated platelet GPIIbIIIa receptors positively influences the short- and longterm outcome after stenting, by reducing the risk of subacute thrombosis, by optimising the blood flow in the tributary microcirculation, and by possibly preventing the development of restenosis.24 Also patients affected by diabetes mellitus, advanced in years, or in whom particularly complex coronary lesions have been detected, seem to benefit from stent usage, more than by PTCA alone.25±27 In 1977, it was discovered that systemic delivery of heparin suppresses the formation of intimal hyperplasia after injury of carotid arteries in rats.28 Later studies revealed that heparin inhibits proliferation and migration of smooth muscle cells probably by interfering with growth factors and independently from its anticoagulant properties.14,29 Systemic administration of heparin has yielded conflicting results with respect to its effect on intimal hyperplasia in experimental vein grafts.30,31 The late 1980s saw the move away from systemic to local therapies. To minimise systemic effects, delivery of a therapeutic agent locally at the time of the operation would be a logical strategy.32 The current coatings applied to stents are mainly directed to counteract intimal hyperplasia by their inhibiting effect on proliferation. The thrombogenicity of these coatings is not studied in detail, but a high incidence of thrombotic occlusions after implantation of stents coated with drugs that affect the vascular tone indicates a potential serious side effect of these inhibitors of proliferation.

15.1.3 Extracorporeal circulation In 2003 almost 700,000 heart operations worldwide were performed with the assistance of extracorporeal circulation (ECC). With respect to the age and physical condition of patients, a shift toward older adults on one hand and younger children (up to 50% infants, half of them new-borns) on the other was observed. The possibilities of operating on these patients successfully can be explained by continuous improvement in the operative techniques, as well as perioperative supervision and mechanical circulatory support systems. The optimisation of the surfaces of ECC devices is of increased importance, both during the operation (heart-lung machine) and after operation with systems providing sustained heart support. Next to the technical perfection of all these systems, it is also particularly important to consider the surfaces presented to the circulating blood. Since its first use in 1953, extracorporeal circulation (ECC) during open heart surgery has developed into a routine procedure. Nevertheless, an insufficient

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haemocompatibility of materials used for ECC devices still remains a problem. The contact between blood and the various artificial surfaces of the extracorporeal system leads post-operatively to a post-pump syndrome, which can escalate into a systemic inflammatory response syndrome (SIRS),33 acute lung failure (ARDS: adult respiratory distress syndrome),34 sepsis, or even multiorgan failure (MOF).35 The causes of these syndromes are multi-factorial; mechanical and chemotactic activation and membrane-damage of the blood cells, dysfunction of cellular immune regulation, and activation of the haemostatic system. The materials used for extracorporeal application include a wide spectrum of polymers, in particular polyethylene, polypropylene, polyvinylchloride, polyester, polystyrene, polyurethane, and silicone. Although these products possess the required physical properties, they display more or less the same disadvantage; an incompatibility with blood and tissues. Through contact with the blood, this incompatibility can provoke a pathophysiological response from the organism, similar to that of traumatic shock. As is well known, in adult patients undergoing a bypass grafting procedure the total blood volume comes into contact with about 3 m2 of these non-physiological surfaces for one to several hours. This extensive contact causes a massive activation of the humoral and cellular defence systems. Such side effects of biomaterials are counteracted in part by coating surfaces to obtain improved biocompatibility or by pharmacological inhibition of the enzymes responsible for consecutive activation of the cascade reactions

15.2 Surface interactions of blood After contact of blood with a material various proteins will be deposited within split seconds. The main proteins adhered to a surface are albumin, fibrinogen and immunoglobulin, based on their high concentrations in blood. After the initial adhesion a continuous exchange with free proteins takes place, that reaches equilibrium after approximately two hours. This results in binding of higher molecular weight proteins. The relatively medium molecular weight protein albumin will be exchanged in part for larger proteins. Next to nonspecific protein deposition, some components of the contact system react specifically with negatively charged surfaces. As soon as the blood comes in contact with a negatively charged surface, Factor XIIa fragments are formed. These fragments then initiate the entire contact system. -Factor XIIa converts prekallikrein into its active form, kallikrein, which generates the vasodilator bradykinin.36 The deposition and conformation of some plasma proteins on the artificial surface, such as Factor XII, fibrinogen, and vitronectin are a significant criterion for further thrombogenicity.37 After deposition to some surfaces fibrinogen leads to a strong adhesion of platelets through platelet glycoprotein receptors (GpIIbIIIa), followed by platelet aggregation, and release of

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15.3 The main products involved in coagulation, including the intrinsic and extrinsic pathway. Although the intrinsic pathway has physiologically almost no meaning, it plays an important role in activation by biomaterials. The intrinsic pathway is initiated by the contact system (XII and kallikrein) or by Factor XI. Deposition of leucocytes results among others in tissue factor activation, whereas activated platelet membranes contribute to coagulation by complex formation of factors IX, VIII and X, V.

procoagulant contents from platelets. Additionally, contact activation induces activation of the coagulation cascade (Fig. 15.3). Fibrin clots and thrombi are cleared by fibrinolysis, the enzymatic process of fibrin fragmentation and platelet release from binding to fibrinogen. The fibrinolytic enzyme plasmin can be formed through either the release of endothelial tissue plasminogen activator (t-PA) or by kallikrein-activated urokinase. The inflammatory reaction is initiated by complement activation. C3b, which is present in small amounts in blood, after adhesion to a negatively charged surface forms a C3 convertase (C3 cleaving enzyme) when not immediately degraded by complement inhibitors. Since foreign surfaces lack complement inhibiting capacity, the complement convertases will amplify the complement reaction by cleavage of new C3 molecules, resulting in C3b generation and its deposition onto the surface. Simultaneously, the smaller C3a fragment is released in plasma and this fragment is often used as a marker of complement activation. Thus, an exponential activation of the complement system takes place after recruitment of the other complement factors (Fig. 15.4). Similarly to C3 convertase C5 convertase will also be formed, with cleavage capacity for C5 into C5a and C5b conversion. The complement has lytic effects on target cells by its end stage components C5±9 (terminal complement complex) and therefore may become harmful for

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15.4 The alternative pathway of the complement system reacts independently of other immune factors on any foreign surface by deposition of factor C3b and subsequent recruitment of other factors, such as Bb. The C3 convertase cleaves new C3 molecules to form C3b and C3a, which enhances deposition of C3b and causes signalling of leucocytes by release of C3a and C5a. Finally, the terminal complement complex, composed of all factors C5b±C9, is formed. This complex may cause host cell lysis.

the patient in contact with an activating device. Moreover, most of the deleterious effects of complement activation are related to the recruitment and activation of leucocytes, such as granulocytes and monocytes. Granulocytes show an upregulation of the adhesion molecules CD 11b and CD 18 with increased adhesion to the surface, release of elastase and superoxide generation, i.e., further propagation of the inflammatory response.38,39

15.2.1 Leucocytes Leucocytes release a number of inflammatory products including chemotactic factors, growth factors, and complement components. A second mechanism involves the production of lysosomal degradation enzymes. Activated leucocytes elaborate several potent proteases capable of degrading collagen and other structural extracellular matrix and extracellular components, for example, basement membranes. Heparanases can remove heparan sulphate proteoglycans from the cell surface and diminish their inhibition on cell proliferation.18 Lastly, leucocytes may also act at sites of endothelial injury through the production of oxygen free radicals. Granulocytes can produce oxygen free radicals capable of injuring remaining viable endothelium leading to an ongoing stimulation of inflammatory injury.14,18

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Inflammatory processes related to biomaterials have been studied extensively in patients undergoing extracorporeal circulation (heart-lung machine or dialysis). During surgery and the early post-operative stage, the extent of the inflammatory response is associated with clinical symptoms such as fever, bleeding and in severe cases organ failure.1,40 This is more pronounced after use of the heart-lung machine, due to its large surface area and corresponding massive activation of the complement system. More recently its was found that use of the heart-lung machine causes biphasic complement activation. The first phase occurs during operation and directly results from the interaction of blood with the extracorporeal circuit. The second phase occurs post-operatively and is characterised by increasing levels of acute phase proteins such as secretory phospholipase A2 (sPLA2) and C-reactive protein (CRP) that contribute to complement activation.45 Conversely, the inflammatory reaction during CPB may contribute to the post-operative generation of sPLA2 and CRP41 and to post-operative morbidity, although it also functions to promote phagocytosis of injured cells and tissue debris.42 Coatings of the extracorporeal circuits have improved biocompatibility, resulting in reduced complement activation and reduced activation of leucocytes during and after bypass surgery.43±46

15.2.2 Platelets Endothelial denudation exposes the subendothelial matrix and leads to platelet adhesion and aggregation. The subendothelium is completely covered by platelets immediately after denudation. Platelet adhesion requires the interaction platelet receptor Gp1b, plasma von Willebrand factor and fibronectin. Platelet aggregation requires fibronectin, von Willebrand factor or vibronectin, and most often platelet receptor GpIIbIIIa. The adhered platelets release adenosine diphosphate and activate the arachidonic acid synthesis pathway to produce thromboxane A2.47 Thromboxane A2 is a potent chemo-attractant and smooth muscle cell mitogen and leads to further platelet recruitment.48 Once activated, platelets release constituents of their granules. Upon activation and during apoptosis platelets and other cells bud off small parts of their plasma membrane, the so-called microparticles (MP). Extensive in-vitro studies have been reported on platelet-derived microparticles (PMP).49±52 When platelets are stimulated in vitro with agonists such as a combination of -thrombin and collagen or the complement complex C5b±9, they release large numbers of PMP. PMP possess `platelet factor 3 activity', i.e., they facilitate coagulation via exposure of negatively charged phospholipids, thereby providing binding sites for activated coagulation factors V (factor Va), VIIIa, IXa, XIa53,54 and enabling the formation of tenase- and prothrombinase complexes (Fig. 15.3).55 Increased numbers of PMP have been reported in the circulation of patients undergoing cardiopulmonary bypass (CPB), and those suffering from acute coronary ischaemia,56-58 and after plasmapheresis,59 which

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has been associated with a thromboembolic tendency. Also in the pericardial fluid of patients undergoing CPB surgery elevated numbers of PMP have been found.60

15.3 Role of blood cells during flow: rolling of cells, effect of concentration of erythrocytes, expression of adhesive cell receptors The rheological properties of blood influence adhesion of platelets, capture and rolling adhesion of leukocytes as well as their margination in the bloodstream.61 Increasing erythrocyte aggregation correlates with increasing leukocyte adhesion and with more slow-flowing leukocytes near the wall. Thus flowing erythrocytes promote leukocyte adhesion, either by causing margination of leukoctes or by initiating and stabilising the attachment that follows. During cardiopulmonary bypass (CPB) a number of non-physiological events take place including haemodilution, hypothermia, and non-pulsatile flow. As a consequence of these events rheology changes and blood flow may be stimulated to shunt from a less favoured organ to preserve a more vital organ.62 At the onset of CPB the prime solution of the extracorporeal circuitry (ECC) mixes with the patient's blood volume. Dilution of corpuscular blood cells, plasma proteins, a reduction of the colloid osmotic pressure (COP) and oxygen transport capacity takes place. To overcome reduced viscocity by haemodilution, to meet the metabolic demands and protect the tissue from ischemia, moderate hypothermia (28 C) is used. In many institutions the target percentage of red blood cell mass in whole blood (haematocrit) during CPB is 22% to reduce transfusion requirements during open heart surgery.63 If possible, autologous blood is predonated and replaced by an equal volume of an artificial colloid to reach the predicted haematocrit. However, in a study where prime volume was reduced during CPB, the haematocrit in the reduced prime group was significantly higher than in the full prime group, resulting in a reduction in allogeneic blood use.64 Moreover, markers of organ ischaemia are reduced when the haematocrit is kept to a more physiological level. During CPB, specifically in non-pulsatile mode, splanchnic vasoconstriction and hypo-perfusion occur, leading to intestinal mucosal ischaemia, and subsequently to the release of endotoxins from the gut into the circulation.65,66 Circulating endotoxins seem to be one of the most potent factors in CPB that contribute to post-operative morbidity by producing an inflammatory reaction.67 Although the more vital organs such as the kidneys show no adverse effects in the course of CPB,68 during rewarming inadequate perfusion and subsequently potential damage to the superficial cortex may occur.69 Ranucci,70 clearly demonstrated that a haematocrit value below 25% is the primary risk factor in developing severe renal dysfunction. It is suggested that an optimal haematocrit during CPB is not necessarily a lower haematocrit, because it should lead to

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sufficient overall tissue oxygenation and reduce any preferential blood flows to other vital organs, whilst at the same time limiting the requirements for blood transfusion.

15.4 Biomaterial surface characteristics in relation to haemocompatibility and clinical applications A very important requirement for biomaterials used for temporary support of organs or as permanent implants in the human body is minimal generation of thrombosis. Adhesion and activation of platelets to biomaterials surfaces is an important step in thrombosis and is governed, in part, by surface energy and wettability of the biomaterial surface.71 Prior to adhesion of platelets, plasma proteins like fibrinogen and fibronectin adsorb,72 and the composition of the adsorbed plasma proteins relates to the wettability of the biomaterial surface.73,74 Adhesion can be controlled by adjusting the surface properties ± especially surface energy ± of the material involved. Long-term implantation of totally artificial hearts is one of the most compelling proofs of the bioengineering utility of surface energy modification to minimise biological adhesion. These pumps, and the related intra-aortic balloons and left ventricular assist devices, do not accumulate blood clots or thrombotic masses during their contact with blood.

15.4.1 Wettability The effects of wettability on deposition of thrombotic material can be studied by means of wettability gradients. These are made by a gradual change in chemistry along their length, resulting in a varying wettability. Therefore, they are excellent tools for studying the biological effect of this property in one systematic experiment. These wettability gradients have been employed to study various biological phenomena, including protein adsorption,75,76 and cellular adhesion.77±79 Recently, a new, simple method to prepare wettability gradients on polymers by means of glow discharge by partly shielding the material with an aluminum cover has been published.80 Such prepared wettability gradients on polyethylene extended over 4 to 5 cm, and their steepness could be controlled by adjusting the height of the aluminum cover above the polyethylene surface and the duration of the treatment. Consequently, shielded gas plasma produced wettability gradients on polyethylene are very suitable to study biological interactions, because they extend over appreciable lengths and the gradients are relatively stable. It has been hypothesised, that shielded gas plasma-produced wettability gradients are a result of hydrophilic groups on a hydrophobic base.81 Albumin, fibrinogen and immunoglobulin G are the most prevalent proteins in blood plasma82 and their adsorption along the length of a shielded gas plasma-

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produced wettability gradient on polyethylene increased with the distance from the hydrophobic end. Since most implanted devices are in contact with blood under flow conditions, it is also relevant to study the effect of wettability of a device surface on the adhesion and activation of platelets under conditions of flow, especially because high flow may trigger the adhesion and aggregation of platelets.83,84 In vitro, shear force induced platelet activation and adhesion to collagen occurs within 2 s, with half of the number of reacting platelets adhering within 240 ms.85 This is important for efficient haemostasis under flow conditions and in the contact of blood with medical devices. Generally, more platelets adhere to the hydrophilic than to the hydrophobic end of a gradient, while flow promotes platelet adhesion evidently through increased convective mass transport.86 Moreover, attachment is known to be stimulated by shear stress, which causes haemostasis under arterial flow conditions.87±90 A moderate flow and shear stress (0.8 N/m2) generated the most pronounced difference in platelet adhesion along the gradient surface. However, when the flow was further increased to simulate the conditions of coronary arteries at 3.2 N/m2, platelet numbers at the hydrophilic end were significantly reduced as compared with the hydrophobic end. These results strongly suggest detachment of platelets from hydrophilic surfaces. Within 15 min at high shear force platelet deposition on hydrophilic surfaces will be limited, probably by detachment after initial adhesion. Such effects can be explained by the small contact area of platelets with hydrophilic surfaces.78 Furthermore, the platelets attached to hydrophilic surfaces remain spherical and extend deeper in the blood flow, thereby experiencing higher shear forces. In contrast, platelets on hydrophobic materials can withstand high shear forces due to strong contact and complete spreading of platelets. When examined with scanning electron microscopy, the platelets on the hydrophobic end of a gradient surface were indeed more extended like a pancake than on a hydrophilic end.91 Obviously, the gradient surfaces at the hydrophobic end reacted similarly with platelets under a high shear stress of 3.2 N/m2 as under a low shear stress of 0.8 N/m2, which indicates a relatively strong binding of platelets on the hydrophobic surface. It can be concluded that under conditions of arterial flow, especially at the hydrophilic end of the gradient surface, fewer platelets adhere after 15 min due to detachment.92 It is hypothesised that hydrophilic device surfaces exposed to flowing blood in the human body and under high shear conditions, are less likely to accumulate platelets than hydrophobic surfaces.

15.4.2 Roughness The influence of biomaterial roughness on thrombogenicity is not clear, since various studies show apparently conflicting results. Bailly compared angiographic catheters with different surface roughness by their tendency to become occluded. They concluded that the most thrombogenic material was the

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smoothest, whereas surface chemistry (polyethylene versus polyamide) contributed to a lesser extent to thrombogenicity.93 Zingg et al. found that increased roughness caused a decrease in platelet adhesion on hydrophilic surfaces and an increase on hydrophobic surfaces. These results were obtained when flow conditions were applied. During static test conditions no differences between smooth and rough surfaces were found.94,95 One explanation for different observations is the higher extent of thrombogenicity at smooth surfaces, whereas the degree of thrombus adhesion is higher at rough surfaces.96 In a more detailed study it was observed that roughness due to titanium crystals appeared to initiate more activation of the clotting cascade, but less platelet adhesion.97 These different effects of two important factors of thrombus formation in conjunction with the variability induced by various flow and shear stress conditions and wettability may explain the conflicting results regarding the thrombogenicity of biomaterials.

15.5 Haemocompatibility of metals, ceramics and polymers Due to their mechanical and radio-opaque properties metals are frequently used for manufacturing implant devices and as part of devices used for invasive procedures for diagnostic and therapeutic purposes. Often, these implants are in direct contact with blood, e.g., as stents, heart valves and catheter tips. The blood compatibility data of metals are relatively scarce, which is possibly due to the historically accepted application of metals as medical devices, or to the indispensable physical characteristics of metals. In reviewing the literature, some frequently used metals even appear to have prothrombotic properties which, if alternatives were available, would lead to refusal of its use in bloodcontacting devices or implants. No thorough comparisons between metals can be made, since the reported studies have evaluated the blood compatibility of different materials and no consistent reference materials were included. Furthermore, most studies report only limited blood compatibility tests. Apparently, the possible induction of an inflammatory reaction, initiated by complement activation or granulocyte activation is not frequently tested, although it is an important contributor to intimal hyperplasia.98 It can be concluded that the more noble metals appear less blood compatible than the oxidised titanium and aluminium metals (ceramics) and silicon carbon products. Most bare metals had a poor blood compatibility in direct comparison to polymers, which was most often tested after polymer coating of the metals, indicating that the mechanical properties of metals are still considered essential for stent or valve construction. A thorough evaluation of the blood compatibility of metals is warranted to quantify their thrombotic and inflammatory properties. Some of the most frequently used metals are now discussed.

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15.5.1 Stainless steel Stainless steel (316L) is the most commonly used metal for endovascular devices. Its mechanical properties significantly contribute to its applicability, but the blood compatibility results also appear better than those of some other metals. For instance, stainless steel stents are more blood compatible than tantalum stents.99 However, stainless steel can also be further optimised, since several studies showed that polymer coating of stainless steel stents reduced deposition of platelets and thrombus mass by more than 60%.100,101 The reported reduction by stainless steel of the partial thromboplastin time by 50% or more, indicating activation of the clotting system, would be considered unacceptable in view of the requirements for newly developed biomaterials.

15.5.2 Tantalum After stainless steel, tantalum was introduced as the metal for the construction of stents. Due to tantalum's high radiopacity, implantation of tantalum stents is greatly facilitated. Initial studies showed similar blood compatibility for tantalum and stainless steel,102 although later studies indicated that stainless steel possesses better blood compatibility.99 Clinical studies indicated that a high incidence of thrombotic complications could occur after tantalum stent implantation if anticoagulation and anti-platelet therapy was insufficient.103 Also, post-stent antithrombotic therapy was required, including both anticoagulants and platelet inhibitors or Ticlopidine plus Aspirin.104 Polymer coating of tantalum stents with polyurethane or parylene reduced the deposition of platelets by 5 to 50% relative to platelet deposition on uncoated stents.105

15.5.3 Titanium In the human body, titanium exists only for a short period of time in its unmodified form, and relevant blood compatibility data are therefore obtained with titanium nitride or titanium oxide. Titanium oxide appears to reduce fibrinogen deposition due to its semi-conductive nature. This effect is explained by the similar electronic structures of fibrinogen and titanium.106 In a comparative study with low-temperature isotropic pyrolytic carbon (LTI carbon), not only reduced deposition of fibrin, but also a 50% reduction in microscopically counted platelets was observed with titanium oxide.107 Transvenous inferiorvena-cava filters made of stainless steel, titanium or titanium-nickel all showed approximately 25% early thrombosis in clinical use, measured via ultrasound scanning.108 This incidence of early thrombosis was unexpectedly high, and difficult to reduce with the current devices, since antithrombotic medication is often contra-indicated in patients requiring a vena cava filter. Titanium nitride has been tested for its blood compatibility with regard to

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leukocyte adhesion, and appears to retain no leukocytes.109 Additionally, platelet retention was as low as with silicone elastomer, but comparisons with more blood compatible materials have not been made.110 In-vivo experiments with titanium nitride heart-valves in sheep showed some deposition of fibrin and platelets.111

15.5.4 Nitinol Nickel-titanium alloy (Nitinol) has attracted special attention due to its shape memory function. It must be noted that Nitinol has an outer surface of titanium (oxide), whereas nickel is not exposed to blood. Therefore, blood compatibility characteristics are expected to be rather similar to those of titanium oxide. Based on the hypothesis that a semi-conductor prohibits fibrin and platelet deposition, Nitinol is expected to be thromboresistant, unless its semi-conductive nature is lost in the alloy. In a clinical study with Nitinol intravascular clot filters, the effects on the clotting system and on platelet adhesion were shown to be similar to those induced by stainless steel.112 An experimental study with stented rabbits showed significantly more thrombus formation on stainless steel than on Nitinol.113 However, grafting of polyethylene oxide (PEO) on Nitinol reduced the fibrinogen adsorption by as much as 99%, and significantly reduced platelet adhesion, which once more shows the superior thromboresistant effects of polymers as compared to those of metals like Nitinol.114 Further evidence that Nitinol, too, can only be safely implanted during antithrombotic treatment was provided in experiments that included the use of platelet inhibitors Aspirin and Copidrogel in a porcine stent model. Combined treatment with these inhibitors reduced stent thrombosis by 95±98%.115 An effective coating such as PEO could thus limit the use of systemic treatment by medication.

15.6 Biological surface treatment to improve haemocompatibility Immediate stimulation of platelets during contact with non-albumin coated extracorporeal circuits has been observed by the release of platelet granules. After this first contact no further substantial platelet stimulation occurred. Albumin has been described to inhibit this release reaction of platelets and to inhibit platelet aggregation.116,117 The first pass effect can be reduced by initial adsorption of albumin to the biomaterial surface. Fibrin deposition and platelet GpIIIa receptor binding was also reduced on tubing of an extracorporeal system after pre-coating with albumin. It appeared that a low concentration of albumin was as effective as a high concentration on reducing platelet activation and on reducing deposition onto the tubing. Since albumin priming seems to exert an inhibiting effect only in the first period of extracorporeal circulation, the first

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pass effect on untreated biomaterials seems to exert the most pronounced haemocompatibility problem. However, the initial albumin coating will be replaced by other proteins during prolonged blood contact as a result of the Vroman effect.118,119 A further improvement of the extracorporeal circuit is supposed to be related to more irreversible surface treatment. Dependent on the treatment, the surface may be modified to induce less thrombogenic or inflammatory reactions. Heparin, cell membrane phospolipids, and block copolymer coatings are often used on a number of blood contacting devices.120

15.6.1 Heparin coating The attempt to coat artificial surfaces with heparin, a natural anticoagulant, illustrates the first step to improve haemocompatibility. Gott reported in 1963 on heparin coating of synthetic materials, which had been pre-treated with colloidal graphite.121 Larm122 presented a heparin-coating method in 1983, which is still the most stable and most effective for long-term use. Their method involving covalent binding by the technique of end-point immobilisation, did not adversely affect the heparin active structure and thus produced a bioactive surface. After the first heparin-coated extracorporeal circuits became available in the last half of the 1980s, its haemocompatibility was shown in in vitro systems, animal models, and patient studies.123±125 It can be concluded that heparin coated circuits can cause a reduction of activation of the contact phase, complement system activation, inflammation, and pulmonary complications.126±128 The reduced thrombogenicity of the heparin-coated surfaces was thought to be attributable to the inhibition of thrombin by catalysing the binding to antithrombin III. However, more recent data show that the advantage of heparin coating lies much more in the reduced, or selective, adhesion of plasma proteins. This leads to a faster formation of a blood-friendly secondary layer and prevents a further denaturation and hence activation of the adhered proteins and blood cells. To minimise adsorption of proteins and attachment of cells, next to the heparin effect, Trillium Bio-passive Surface has been developed. This technique works with water-soluble synthetic polymers that are immobilised in two superficial layers. The first polymer is a primer and is based on polyethyleneimine that is modified hydrophobically to allow for strong binding to artificial materials of the medical device. The second layer, containing sulfonate groups, polyethylene-oxide chains, and heparin, is covalently bound to the primer and results in an insoluble surface coating. This surface prevents the adhesion of blood cells and plasma proteins during contact with blood by its negative charge, heparin and PEO chains.129±132 The first experiments with this technique indicated that the treated synthetic materials prevented the activation of the complement system, the contact system

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as well as platelet and leukocyte activation. The TBS coating intervened in the initial phases of the blood-material interactions and thereby prevented further activation at a very early stage in the cascade reaction. ECC systems with this coating did not, however, significantly improve the clinical outcome of systemic heparinised adult patients undergoing cardiac surgery.133

15.6.2 Phosphorylcholine coating Phospholipids that mimic natural cell membranes are currently used as a coating substance. Phosphorylcholine-containing lipids dominate in the outer membrane of the cell membrane bilayer and these appear to possess strong antithrombotic properties.134 One has succeeded in coupling synthetic methacryloylphosphorylcholine/lauryl-methacrylate copolymers to metal and synthetic surfaces. The term `biomembrane mimicry' arose for phosphorylcholine-coated foreign surfaces.135 In-vitro experiments and animal tests have shown that phosphorylcholine-coated artificial polymers possess outstanding thrombogenic resistance and display only minimal adhesion of plasma proteins and platelets.136,137 This coating technique has been offered among others for contact lenses, stents and extracorporeal circulation devices. Since coagulation in infants is more delicate than in adults by the reduced availability of inhibitors, this antithrombogenic coating was anticipated to be most profitable for paediatric cardiopulmonary bypass. Although the literature shows an improved biocompatibility in adult surgery when using coatings,138 the thrombogenic and inflammatory response is usally mild in routine adult surgery which makes it difficult to demonstrate differences in post-operative clinical response. Small infants are much more vulnerable to the adverse effects of cardiopulmonary bypass due to the relatively high priming volume and large blood-foreign material surface in combination with the immaturity of several organ systems. A clinical study on small infants showed a reduced thrombogenic and inflammatory response after the use of phosphorylcholine coating, by reduced progression of beta thromboglobulin release and thromboxane production, both related to platelet activation, and by reduced complement activation. While the surface characteristics improved, the coating did not affect the gas transfer properties of the hollow fibre membranes. The characteristic feature of biological membranes is their functional and compositional lipid asymmetry, which has been described in several cell types and is thought to stem from the requirement for biological membranes to have asymmetric protein distributions across the bilayer. In all of the cells for which lipid compositional asymmetry has been described, negatively charged phospholipids are found predominantly on the inner cytoplasmatic side of the membrane, while the neutral zwitterionic phosphorylcholine-containing antithrombotic lipids predominate in the outer membrane leaflet. Negatively charged phospholipids are thrombogenic and it has been proposed that this

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membrane asymmetry may serve the biological purpose in the maintenance of the delicate balance between haemostasis and thrombosis. In-vitro experiments, in which various phospholipid coatings were applied to surfaces, showed a very high procoagulant activity of negatively charged phospholipids. This is in contrast to phosphorylcholine-containing surfaces that were not active in coagulation tests.139,140 However, inhibition of activation of the clotting system was not observed, which may indicate a merely passive effect of the phosphorylcholine coating towards the clotting system. Previously, heparin coating has been evaluated in paediatric CPB.141,142 As in adult CPB heparin coating reduced complement activation. Surprisingly, also the phosphorylcholine coating appeared to generate less complement activation than the uncoated systems. In-vitro experiments showed decreasing complement activation with increasing surface phoshorylcholine mole fractions,143 suggesting that phosporylcholine is responsible for the reduction. The working mechanism is probably related to lesser activation of the complement protein C5144 and the inhibition of monocyte and macrophage adhesion.145

15.6.3 SMA coating Surface modifying additives (SMA), are mixed with the initial synthetic materials in the production phase, and this technique is therefore not a coating in the usual sense. The copolymer distributes itself in the synthetic materials during the polymerisation process and due to its charge characteristics, moves to the surface of the basis material as it cools. Thus, a new surface of primarily SMA forms. The microscopic structure of the surface of alternating hydrophilic and hydrophobic regions carries a zero net charge, thereby reducing platelet and leukocyte deposition. Tsai et al.146 could prove that SMA surfaces decreased coagulation activation and significantly reduced contact phase and complement activation. Gu et al.147 found better platelet protection in clinical CPB by using SMA treated devices. However, larger clinical studies on routine cardiopulmonary bypass patients showed only minor clinical benefit of SMA treated devices.

15.6.4 PMEA coating Poly-2-methoxyethylacrylate (PMEA) is a hydrophilic polymer coating that minimises the adsorption and denaturation of proteins and blood cells. In various animal and clinical studies, this coating has been proven to reduce blood activation during extracorporeal circulation. When compared with uncoated oxygenators, PMEA-coated oxygenators exhibited increased thrombus resistance with lower inlet pressure and lower thrombocyte consumption. Plasma bradykinin levels and the percentages of activated monocytes in PMEA-coated circuits were significantly lower than those in uncoated circuits during CPB. The

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amount of protein adsorbed on PMEA-coated circuits was significantly lower than that on uncoated circuits (0.30 g/cm2 versus 3.42 /cm2). Almost no IgG, IgM, or C3c/d was detected in proteins adsorbed to the PMEA-coated circuits although these proteins were clearly detected on the surfaces of uncoated circuits.148 A clinical study showed no significant differences between heparin-coated and PMEA-coated groups in the plasma concentrations of inflammatory markers, or markers of clotting. Clinical variables did not differ significantly between the groups. It was concluded that PMEA-coated CPB circuits are as biocompatible as heparin-coated CPB circuits and prevent post-operative organ dysfunction in patients undergoing elective coronary artery bypass grafting with CPB.149 The cost-effectiveness ratio seems favourable for PMEA-coated circuits.150

15.7 ISO 10993 requirements for testing of medical devices: simulation of clinical application including flow, blood composition, anticoagulants In December 2002 the revised ISO 10993-Part 4 standard (Biological evaluation of medical devices ± Selection of tests for interactions with blood) was published.151 Ten years of discussions, conferences, writing and implementation of new insights preceded the revision of this standard that deals with the reaction of blood to medical devices. The standard is applicable to external communicating devices, either with an indirect blood path (e.g. blood collection devices, storage systems) or in direct contact with circulating blood (e.g. catheters, extracorporeal circulation systems), and implant devices (stents, heart valves, grafts). Testing should be performed for five categories, based on primary processes: thrombosis, coagulation, platelets, haematology and complement. In this system all relevant aspects of blood activation are taken into consideration, but, and this is most important, testing should simulate clinical conditions as much as possible. The increased use of medical devices for temporary use or implant in the blood circulation has resulted in increased demand for evaluation of complications brought about by these devices. One important and extremely relevant aspect of testing of medical devices is the condition of blood exposure to the device. Often, blood with clinically inapplicable anticoagulants and under static conditions was incubated with the test device.152±154 Currently anticoagulation and flow conditions must be as similar as possible to the clinical application to achieve relevant test results. Thus, most devices must be tested with heparinised blood under circulating conditions. For some devices, such as stents and catheters this implies high flow through or around the device to obtain relevant shear stress conditions. The major differences observed between cell interaction

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under static and flow conditions has made clear that whole-blood flow models are required for testing haemocompatibility inasmuch as the test device will be used clinically in the blood circulation. Flow models for testing may consist of animal models or in-vitro test systems. Animal models have the disadvantage of being expensive, time consuming and insensitive due to overwhelming shortterm effects of tissue damage. Animals, particularly, are being used to test haemocompatibility of implants. Clearly, human volunteers cannot be used for this purpose. Animal models include the effect of tissue damage by operation and the important antithrombotic effects of endothelial cells. On the other hand the extent of haemocompatibility is obscured by these tissue effects. Moreover, it has been shown that the composition of blood differs considerably between various species, which leads to over- or under-estimation of human blood reactions to biomaterials.155,156 The use of human blood is therefore more relevant to the interpretation of results and offers a more detailed array of test methods, since most available methods are based on human blood components. The use of human blood requires a proper in-vitro circulation model, which is discussed in the next section.

15.8 Test models: static, low flow, arterial flow, pulsatile/laminar flow A key determinant of blood activation and adhesion of cells is wall shear stress; the force exerted by the flow per surface area. In a cylindrical tube this property is easily calculated from  ˆ 32Q=…D3 † where  is the shear stress, Q is the flow rate,  is the viscosity, and D is the tube diameter.157 In configurations that differ from the cylindrical tube, such as just after a bifurcation, wall shear stress has much larger values locally, than at the opposite site.158,159 Growth of intimal thickness is often observed at locations with low shear stress.157,160 Wall shear stress in the normal circulation is rather constant when the equation is applied to blood vessels of various sizes.157 Values are found in the range 10±20 dynes/cm2 (1±2 Pa). Furthermore, blood vessels adapt their diameter as much as possible towards a constant value for shear stress.161 Also, when blood is in contact with biomaterial surfaces, fluid mechanics, and especially the shear stress, have a strong influence on the damage of red cells and platelets. Red cell damage may occur at high shear stress.162,163 Platelets are more easily damaged by shear stress.162,164 Platelet damage is not only influenced by the maximum shear, but also by the duration of the shear force. Only for very short exposure times are platelets able to withstand higher shear stress than red cells.164 From a fluid mechanical point of view, differences in flow situations may therefore lead to different problems with blood. Artificial heart valves may cause problems for red cells due to short duration high shear,165 whereas stents

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in the coronary arteries induce intimal growth at locations of relatively low shear,166 which may be caused by platelet activation in high shear. Neointima formation in stents has been shown to be related to wall shear stress as well.166 In tubing used during dialysis, the high shear rate at the needle may lead to problems for red cells,167 but it should not be disregarded that the wall shear stress of the tubing is the most critical issue for platelet activation. Heart valves, extracorporeal systems and vascular grafts or stents induce relatively high shear forces that may result in platelet activation. Shear stress is a natural activator of platelets. The shear-induced pathway appears to be one of the major pathways of platelet induced haemostasis and thrombosis.90 The sequence of the shear-induced pathway is the binding of von Willebrand Factor (vWF) to the platelet glycoprotein Ib (GpIb) receptor, the expression of activated GpIIbIIIa receptors and release of platelet vWF. Finally, vWF binds to GpIIbIIIa, leading to irreversible adhesion. High shear stress of 120 dynes/cm2 induces immediate expression of GpIIbIIIa receptors and release platelet vWF multimers.168 However, in the presence of platelet activators, such as epinephrine and ADP, shear stresses of 60 dynes/cm2 may synergistically result in platelet aggregation.169 Fibrinogen appears to mediate platelet aggregation efficiently at low shear rates, but not at high shear rates.170 Moreover, resting platelets do not adhere efficiently to fibrinogen-coated surfaces; activation by ADP is required.171 For understanding the effects of platelets during use of biomaterial implants or extracorporeal systems these observations are important. A biomaterial surface is initially mainly coated with albumin, immunoglobulins and fibrinogen from plasma. Adhesion of platelets to fibrinogen onto these surfaces is not as irreversible as to collagen coated with vWF multimers of a damaged blood vessel. Additionally, platelet activators significantly support, or are even required for platelet adhesion to fibrinogen, which means that concomitant tissue or blood damage, such as surgical trauma, occlusive pumps, or poorly designed heart valves, synergistically contribute to platelet activation under shear stress. The wettability of biomaterials relates to the extent of hydrophobicity and hydrophilicity. An increased wettability on a polymer gradient was related to an increased amount of oxygen incorporated in the material surface85 and appeared to correlate with increased protein adhesion, activation of the coagulation system, and increased platelet adhesion.96 A classical in-vitro test model is the Chandler loop,172 which consists of a closed tubing partly filled with air, which circulates the device constantly with an airliquid interface. This method may induce artefacts due to the major forces applied to blood elements and protein denaturation at the air-liquid interface.173±176 Thus, instead of the Chandler a small roller pump closed-loop system was used in the past. This model appeared effective for short-term circulation.177±179 An initial experimental blood circulation model with a roller pump, already refined in previous studies,180 appeared efficient, reliable and cost-effective in

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assessing the haemocompatibility of stents, possibly before their clinical utilisation. However, blood damage induced by the pump limited the exposure of the test object to circulating blood. Since metal stents are commonly of a thrombotic nature, experiments for 15 minutes yielded sufficient information to compare stents, but a less traumatic circulation system is required for testing of low thrombotic materials. The Chandler model induces less blood damage than the roller pump, but has the major disadvantage of continuous blood-air contact and the limitation of blood flow due to the requirement to keep air at the top of the circuit. Since improvement of the model by minimising blood damage may increase sensitivity and permits prolonged blood exposure,181 we constructed a simple mechanical device without air and without a pump to reduce blood damage and activation by the device. Moreover, the new device provided pulsatile flow at a frequency similar to the arterial circulation. This haemobile was compared with the Chandler and roller pump model for intrinsic blood damage and finally for testing of thrombosis induced by arterial stents. Fast screening of the thrombogenicity of stents, catheters, vascular grafts and other small medical devices is now possible. The adjustable flow and shear in the haemobile renders it a model that allows standardised testing of these devices at the cost of low intrinsic blood damage. A limitation of in-vitro models is mainly represented by the absence of an endothelial layer in the circulating system. Throughout the release of (anti)thrombotic components and the expression of adhesion molecules, endothelium has a major role in mediating the interplay between the injured vessel wall and blood cells after coronary stenting,182,183 and lack of this character can somehow alter the likelihood of our experimental representation. Nevertheless, all the other elements depicting the blood-stent phase boundary scene are present in the in-vitro model.

15.9 Conclusion In spite of all the technical improvements made to improve the haemocompatibility of ECC components, a noticeable activation of plasma proteins and corpuscular blood components still exists. The long-range aim remains the creation of an optimally haemocompatible surface (endothelium-like), which blood would no longer recognise as unphysiological and hence would not induce humoral and cellular defences as well as rejection mechanisms against it. With the results of numerous studies on the haemocompatibility of extracorporeal circulation under consideration, one can definitely work on the assumption that heparin-coated devices retard the activation of cellular and humoral mechanisms. Above all, the reduction of a general inflammatory response initiated by extracorporeal circulation represents an important application for heparin-coated devices. Above and beyond the long-term applications, routine heart operations have also markedly begun to utilise heparin-coated

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devices. This trend in improving haemocompatibility with active coatings will assuredly continue in the coming years. Heparin coatings are merely the beginning of improved haemocompatibility for all materials that come into contact with human blood or tissues. Intelligent materials with almost completely physiological surfaces will be at the surgeon's disposal within the next few years. Such materials will be able to mimic endothelial functions and respond to individual changes in the patient's status with the controlled release of adequate pharmacological substances. However, control of the quality of patient treatment and a more individual approach is essential to improve the quality of treatment with medical devices. On-line monitoring of organ damage (point of care testing) is essential in order to apply optimised conditions. The currently used markers of organ dysfunction are too non-specific or too much dependent on an advanced stage of deterioration of the organ. Therefore, newer, sensitive organ damage markers for early organ damage will probably be introduced in routine practice.184±186 Bare metal parts exposed to blood should be avoided, either by coating, oxidation or replacement by haemocompatible polymers. Further, the blood composition in terms of haematocrit and plasma expanders should be optimalised, to ensure proper oxygen delivery to all organs. A new important item is compliance with magnetic resonance imaging (MRI) techniques, which will be a common method of screening for malignancies, inflammatory sites and infarctions. Stainless steel and nickel could particularly raise problems in MRI scans. Finally, thorough haemocompatibility testing should have a prominent place in the certification of blood contacting medical devices, since a poor haemocompatibility has long-lasting negative effects on the whole body through blood transport of activation products and on functional recovery at the site of the implant.

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16 Cell guidance through surface cues A K V O G T - E I S E L E , Max-Planck Institute for Polymer Research, È U S S E R , Institute for Thin Films Germany, A O F F E N H A and Interfaces, Research Centre JuÈlich, Germany and W K N O L L , Max-Planck Institute for Polymer Research, Germany

16.1 Introduction The guidance of cellular adhesion, migration and network formation in vitro by surface associated cues is of great interest for a number of issues both in basic research and in biotechnological applications: Firstly, processes influenced by cues offered on the substrate, e.g., cell adhesion, migration or morphological alterations can be investigated. This is usually done on surfaces exposing patterns of biologically relevant molecules, such as growth factors or proteins of the extracellular matrix. If cells are seeded, e.g., onto surfaces containing alternating stripes of two different matrix proteins, cellular behavior in choice situations can be studied, monitoring attraction, repulsion or morphological changes along the pattern axis. Secondly, the ability to guide cells and the outgrowth of their processes by surface associated cues is useful for the design and the direction of multicellular assemblies with experimentally defined geometries. Examples of such assemblies are orientated vascular tubes formed by endothelial cells or neuronal networks that interconnect along the geometry of an underlying micropattern. This definition of network architecture can be achieved by presenting a surface composed of adhesive and antiadhesive areas, such that cells are forced to adhere to and connect along the permissive regions, while avoiding the antiadhesive background. Patterned neuronal networks allow the study of signal transduction and signal processing in networks of reduced complexity (since only a few pathways are allowed for cellular connectivity) and of defined geometry (as the connectivity pattern is experimentally determined). Thirdly, the experimental manipulation of cellular outgrowth and connectivity on an artificial surface is a promising tool in biotechnology. One potential application lies in the design of implants and prostheses containing grafted cells, whose interaction with the surrounding tissue of the patient may be controlled by providing defined pathways and interfaces for circuit formation. In addition, cell patterning in combination with microfluidic devices offers attractive possibilities for the creation of novel biosensors. It allows locating

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single cells to defined sites on the surface, positioning them on areas exposed to a stream of applied test substances while others are tethered to unexposed areas. In such arrangements, the impact that drug application to single cells has on the behavior of an entire network can be investigated. A number of different techniques have been applied for surface patterning, which we will briefly review with regard to the guidance of neuronal adhesion and network formation. Methods for surface patterning can be roughly divided into two types. On the one hand, cells have been shown to be guided by topographical cues consisting of shapes and textures in the substrate. On the other hand, differences in the chemical properties of confined surface areas can be employed to direct cell-substrate interactions. Figure 16.1 shows a single neuron following defined pathways of a `cell friendly' pattern while avoiding the repellant background material. However, it is important to bear in mind that the effects of surface chemistry and surface topography are not two entirely separate issues. Differential deposition of culture medium components and/or cell derived matrix proteins on surface features may add (bio)chemical differences to topographical structures, while the chemical modification of a surface also has consequences for its physical qualities such as topography, roughness or elastic properties. In the following subsections, we will briefly compare topographical and chemical patterning methods and give an overview over their history in the application of in vitro neuronal cell patterning. In the second section of this short review, we will summarize general principles for surface modifications by chemical grafting and the involved chemistries, and in section three we discuss

16.1 SEM image of single neurons isolated from the cortex of embryonic rats (18 days gestation) on a micropatterned surface. Patterning was performed by microcontact printing using a blend of extracellular matrix proteins and a background of polystyrene.

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the techniques most commonly used in order to achieve patterned cell growth. Finally, in section four we will focus on the biology of neuronal networks and the impact of different types of surface modifications on cellular physiology as well as recent advances in the geometrical design of neuronal networks.

16.1.1 Topographical patterning The employment of topographical patterning techniques started in the early 1960s (Rosenberg, 1963) and 1970s. Cells were seeded onto planar substrates with etched or scribed grooves to study their alignment with the surface structure. With increasing evidence that curvature was the most important parameter for cell guidance, more groups began to examine the effects of varying groove depth, width, and spacing. In the 1980s, lithography was used to microfabricate grooved surfaces, in particular, by utilizing anisotropic etching of silicon wafers. For further details about concepts, materials, surface structures, and possible cellular and biomolecular mechanisms for topographically patterning we refer to reviews by Curtis and Wilkinson (1997, 1998) and Jung and coworkers (Jung et al., 2001). In the late 1980s and early 1990s, the group around Adam Curtis and Chris Wilkinson in Glasgow started to systematically study the relative effects of groove depth and pitch on cell guidance, applying ultrafine structured quartz and silicon surfaces produced by electron beam lithography (Clark et al., 1990). Later, surfaces containing both chemical and topographical cues were applied in order to determine their relative impact on cell guidance if conflicting directions were imposed by the different systems (Britland et al., 1996). The authors determined the critical depth of topographical features required to override guidance through adhesive stripes of aminosilanes; in addition, synergistic effects were reported if chemical and topographical cues were aligned. A further approach to patterning using topographic cues applies pits and grooves sufficiently deep to trap individual cell bodies and outgrowing neurites (Fig. 16.2). This tactic has been used successfully by the group of Peter Fromherz, who achieved the immobilization of snail neurons on a polyester microstructure which created a `fence' around single cells thus confining them to defined surface spots (Merz, 2002).

16.1.2 Chemical patterning First descriptions of chemical cell patterning were presented in the mid 1960s when Carter and his group discovered that fibroblast adhered preferentially to palladium islands evaporated on a polyacetate surface (Carter, 1965). In 1975, Letourneau applied this method to study the alignment of chick dorsal root ganglion neurons with palladium spots on polymeric substrates (Letourneau, 1975). He could demonstrate that the cells preferentially adhered to the metal

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16.2 Scanning electron micrograph of a topographical structure for the defined formation of neuronal networks. The structure was prepared with SU-8 photoresist. The pit has a diameter of 20 m while the width of the groves is c. 10 m.

regions if these were evaporated onto tissue culture plastic substrates but behaved quite indifferently if the palladium was surrounded by polyornithine. This work showed that differences (`contrast') between adjacent regions were necessary to obtain cell patterning. In 1988, Kleinfeld, Kahler and Hockberger used photolithographic techniques to pattern silicon surfaces with small organic molecules in order to direct the adhesion and outgrowth of neurons (Kleinfeld et al., 1988). This study presented a milestone in the field of neuronal cell patterning, as it systematically analyzed the impact of defined chemical groups on cell attachment. It was found that diamines and triamines, but not monoamines, are supportive of cell attachment while surface bound alkanechains have a strong repulsive effect. Similar to photolithography, organic thin films can be patterned by photochemical reactions. The group of Bruce Wheeler applied selective laser ablation to create high resolution grids of polylysine with varying line width, intersection distance and nodal diameter onto which rat hippocampal neurons were seeded (Corey et al., 1991). Photoablation was also used by other groups to pattern ultrathin polymer layers in order to control the adsorption of proteins and the adhesion and spatial orientation of neuronal cells on surfaces (Bohanon et al., 1996). In the middle of the 1990s the Aebischer group studied the impact of surface bound electrical charges on neuronal attachment and differentiation using patterned polymeric substrates (Valentini et al., 1993; Ranieri et al., 1994). An additional approach to surface patterning by chemical methods is presented by the deposition of organic compounds through plasma polymerization. The method tolerates a variety of different chemistries and permits the creation

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of surface features in the m range if parts of the surface are shadowed by a TEM (transmission electron microscope) mask during the deposition process. The suitability of this method for patterned cell cultures has been demonstrated recently (Bullett, 2001; Mitchell, 2002). Towards the end of the 1980s, evidence accumulated that the extracellular environment profoundly influences cellular differentiation in vitro as well as in vivo (Kleinman et al., 1987). Researchers, therefore, increasingly turned to physiologically occurring matrix proteins rather than synthetic substrates for their cell cultures, the simplest method for surface coating being protein adsorption from solution. In 1986, the Arginine-Glycine-Aspartate (RGD) motif was identified as an epitope shared by many matrix proteins that is central to cell-matrix interactions through binding to integrins, a family of adhesion molecules (Ruoslahti and Pierschbacher, 1986, 1987). Shortly afterwards, a number of additional motifs responsible for cell-matrix interactions were identified (Graf et al., 1987; Kleinman et al., 1990). Following the trend in using physiologically occurring adhesive molecules for cell cultures, both entire proteins and isolated recognition peptides were also employed for micropatterning. In 1987, Hammarback selectively destroyed surface adsorbed laminin in defined positions by UV irradiation through a grid mask. The retention of native and therefore adhesion promoting protein in the areas protected by the mask was demonstrated by the guidance of cells along the narrow stripes of protected laminin (Hammarback, 1985). Several years later, the group of Peter Fromherz successfully applied photolithography to pattern ECM proteins for the guidance of leech neurons in culture (Fromherz et al., 1991; Fromherz and Schaden, 1994). Similarly, adhesion promoting peptides were patterned by laser-lithography (Matsuzawa et al., 2000) or microcontact printing (Scholl et al., 2000). In both studies, cell attachment to the permissive areas and neurite outgrowth along the desired pathways was encountered to a similar extent as on substrates patterned with the entire protein.

16.2 Surface functionalization The deposition of chemical cues to solid substrates by a mere physisorption step, e.g., by adsorption of polymers or proteins from solution, is the simplest way to generate a specific functionality. However, the interaction forces involved are typically weak and, hence, the coatings are prone to detachment, delamination, or desorption, e.g., adhesion proteins bound to surfaces via hydrophobic interactions can be displaced by other (small) hydrophobic or amphiphilic molecules. Even though the many weak Coulombic interactions between the charges along a polyelectrolyte molecule adsorbed to the countercharges at a correspondingly modified surface give a relatively stable layer, a gradual dissolution of the adsorbed polymer by the interference of small ions renders this coating inherently instable.

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The introduction of covalent bonds between groups on the substrate surface and correspondingly reactive moieties of the molecule to be attached provide the coating with the desired long-term stability, even under chemical or physical stress, e.g., against electrolyte attack, temperature gradients, or flow shear-stress conditions. For low molecular mass systems the `grafting-to' approach, i.e., the coupling of a molecule through a reactive endgroup to the surface has proven to be a very versatile strategy for the introduction of a broad range of functionalities or a general modification of surface properties, e.g., the generation of hydrophobic/ hydrophilic interfaces, the control of wettability, the deposition of surface lubrication layers. For chain- or rod-like molecules, the gain in free energy by bond-formation between their reactive head-group and the substrate can overcompensate the loss of entropy of the molecule in going from solution to the surface and, thus, allows for the formation of highly organized, even crystalline surface coatings. Firstly, as demonstrated for long alkyl-chains with multifunctional silane head groups (Sagiv, 1980), this concept of monolayer formation by a self-assembly process has gained a widespread popularity through the chemically more versatile thiol derivatives (Nuzzo and Allara, 1983). While the latter are strongly binding to (noble) metal surfaces and, thus, allow for the stable modification of, e.g., electrode surfaces, the former silyl-derivatives are well-suited to the coupling to oxide surfaces, i.e., for the modification of glasssubstrates or non-metallized oxide gate electrodes of field-effect transistors. If silane- (or thiol-)derivatives also exhibit a reactive end-group, further coupling steps can lead to more complex supramolecular interfacial architectures. An example of such a strategy is given in Fig. 16.3. The surface silanol(SiOH-)groups of a glassy substrate are coated with APTMS (aminopropyltrimethoxysilane) resulting in an amine-terminated surface layer. A so-called hetero-bifunctional cross-linker, i.e., a small molecule with one reactive group specifically designed to couple to amine groups (cf the structure formula in Fig. 16.3) can be attached next, leading to a reactive surface capable of binding ± via its second reactive group ± e.g., the SH-group of a cysteine-terminated peptide or protein. In the example given in Fig. 16.3, the peptide consists of the binding domain of the B2 chain of laminin. Thus, a carpet of peptide motifs is generated which has been shown to be highly attractive for the adhesion of randomly seeded hippocampal neurons (Matsuzawa et al., 1996). This concept offers a number of advantages in addition to being a simple recipe that involves only commercially available coupling reagents. The choice of the linker system can be optimized in terms of its chemical nature and its molecular architecture to the extent that chemical degradation of the final protein layer by a denaturing interaction with the surface can be minimized. Moreover, the aminosilane layer can be easily patterned (cf. below) by photolithographic strategies. By choosing the appropriate UV-wavelength, the Si-O-Si bond can be selectively hydrolyzed again regenerating free silanol groups in the exposed areas for the next self-

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16.3 Peptide immobilization on a surface using a heterobifunctional crosslinker. The silanol-groups of a glass substrate are modified with aminosilane groups, leaving an amine-terminated surface. These amine groups subsequently react with one of the functional groups of a crosslinker molecule, thus immobilizing it on the surface while the second functional group remains free. During the next step, the surface is incubated with a peptide or protein containing a terminal cystein residue. The SH group of the cysteine then forms a covalent bond with the free reactive group of the crosslinker molecule, leaving a carpet of peptide chains on the surface.

assembly step, e.g., with a silane derivative of another end group functionality (Dulcey et al., 1991). The `grafting-to' principle has been shown to work well for low molecular mass molecules. However, in cases where longer, polymeric systems were to be end-grafted, i.e., were to be coupled to the substrate via one of their two end groups, this concept failed for mainly two reasons (RuÈhe and Knoll, 2000). Firstly, the statistical distribution of the two end groups within the random coil of a polymer in solution makes it more and more unlikely that for an increased molecular mass polymer, that one of these will reach the surface and form a required covalent bond with a surface reactive moiety. Secondly, the increase in surface coverage by polymer chains that have already bound to the surface generates an increasing kinetic barrier to the next chain approaching the interface from solution. The osmotic pressure generated by the bound layer slows down the penetration of the endgroups of the new chains and thus prevents them from forming a stable bond with the surface. The `grafting-to' approach therefore leads to only rather thin surface coatings in the range of only a few nanometers.

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If dense polymer brushes are required, a novel concept originally proposed by DeGennes, and experimentally realized by the RuÈhe group can be used. Instead of coupling the whole polymer to the surface, only the small initiator molecule is covalently bound and thus leads to a functional surface coating of high density (Prucker and RuÈhe, 1993). The induction of the polymerization reaction by activating surface-initiators by heat or light in the presence of monomers can then lead to polymer brushes of very high graft densities (down to a chain-chain separation of d0 ˆ 2±3 nm) and with individual chains reaching a molecular mass of 106 Da. This concept allows the use of a wide variety of monomers, thus providing a recipe for the generation of very different surface cues. An additional advantage is the possibility of activating the initiator by photons, thus, enabling the generation of patterned surface coatings using a suitable mask for photo-initiation. This way, not only attractive coatings but also highly cell-repellent areas have been generated on solid surfaces, e.g., by the growth of a patterned brush of only 20 nm poly(styrene), the coated areas on the surface of a transistor chip were completely free from adherent cells after random seeding. Only the protected uncoated areas were covered by a dense monolayer of cells (Prucker et al., 1998).

16.3 Patterning of chemical surface cues 16.3.1 Photolithographic patterning Photolithographic techniques are well established for the mass production of silicon chips with a resolution and alignment precision of different surface areas in the sub-m range. Patterns are created by the partial exposure of photoresist surfaces to UV light through a chromium-quartz mask, followed by a chemical development step during which the exposed areas are etched away. Afterwards, the surface in these regions can be chemically modified, while the resist-covered sections are inaccessible. Subsequent removal of the photoresist with organic solvents leaves well-defined patterns of modified and unmodified surface areas (Fig. 16.4). Although the technique was initially developed for inorganic materials, Kleinfeld and coworkers successfully applied it to pattern thin films of organic molecules (Kleinfeld et al., 1988). One drawback to this system for application in cell patterning is that the solvents and development solutions are chemically relatively aggressive. It was therefore unclear whether the method can also be utilized to pattern biological molecules, which are rather sensitive with respect to the surrounding milieu. In 1993, Clark and coworkers (Clark et al., 1993) found an elegant solution to circumvent this problem by patterning the adhesion molecule laminin indirectly through a photolithographical method. In the initial patterning step, methyl groups were bound to defined surface areas while the remainder were left uncoated. As laminin adsorbs preferentially to hydrophobic surfaces, incubation of the substrates with a laminin solution resulted in a pattern of laminin adsorbed

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16.4 Scheme of the steps involved in surface patterning by photolithography. A quartz substrate onto which a thin layer of photoresist has been applied by spin-coating (a) is irradiated with UV light through a photomask (b). During the following development step, the exposed areas are etched away (c). The surface is then flooded with the desired molecules (e.g. alkyl-trichlorosilanes), which bind to both the photoresist-covered and uncovered areas (d). After removal of the photoresist with organic solvents (e), the surface is incubated with a second type of molecules (e.g. aminosilanes), for which only the previously covered surface areas are available for binding (f). Note that the figure is not drawn to scale, the photoresist layer is  1 m thick, while the alkanes and amines form layers of molecular thickness. Adapted from: Kleinfeld et al., Journal of Neuroscience 8, (1988), 4098±120.

to the methylated surface areas against uncoated quartz. The protein coated tracks were shown to guide the outgrowth of neurites from chick embryo brain neurons, indicating that at least the epitope(s) required for this function had retained their native conformation. Since then, standard photoresist techniques have been adapted to generate micropatterns of proteins on glass by using lift-off and plasma-etching techniques (Tai and Buettner, 1998; Sorribas et al., 2002).

16.3.2 Photochemical patterning Photochemical methods can be used to pattern self-assembled monolayers (SAM) or thin films of organic molecules by exposing the surface to UV light

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through a photomask or a metal mask. Illumination with UV light causes oxidation of the molecules in the exposed areas (for example alkanethiolate oxidize to alkanesulfonate) which alters their chemical properties, e.g., their solubility. Immersion of the patterned substrate in a solution containing another organic molecule results in the modification of the illuminated region by a second monolayer (Dulcey et al., 1991) while the non-illuminated region remains inert. An appropriate choice of molecules for the two layers allows for the creation of patterned surfaces with contrasting adhesive properties suitable for controlling cell adhesion to defined sites. The method has been utilized by Ravenscroft and coworkers who irradiated a thin surface bound film of the cytophilic molecule DETA (trimethoxysilylpropyldiethylenetriamine) through a photolithographic mask, thus creating reactive hydroxyl-groups in the illuminated areas. These reacted with the cytophobic silane 13F (tridecafluoro-1,1,2,2-tetrahydrooctyl-1-dimethylchlorosilane) in a second step which resulted in complementary anti-adhesive surface areas (Ravenscroft et al., 1998).

16.3.3 Plasma micropatterning Plasma polymerization is a method in which gaseous monomers, excited by a plasma, react and precipitate onto two- or three-dimensional substrates as a highly crosslinked polymer layer. Through the right choice of monomers, plasma power input and deposition time, it is possible to control a number of properties of the deposited layer, such as its hydrophilicity, chemical nature, density of functional groups and surface energy. This has been exploited to systematically render surfaces attractive or repulsive to cell attachment. Deposition of n-hexane films, which resulted in cell repulsion, through a transmission electron microscope mask, has been used for the creation of a polymer pattern on permissive tissue culture dishes. Groups in Sheffield and Aberdeen were able to demonstrate the utility of these surfaces by growing neuronal cells on these substrates and show that they adhered selectively to the protected areas of the culture dish (Bullett et al., 2001; Mitchell et al., 2002).

16.3.4 Microcontact printing Microcontact printing (CP) was initially developed by the Whitesides group who used it to print monolayers of alkanethiols onto gold substrates (Kumar and Whitesides, 1993). An elastomeric stamp is employed to create patterns of organic molecules on planar surfaces. The application of this technique to the investigation of cell-substrate interactions has mainly focused on endothelial cell adhesion and control of neuronal process outgrowth for the creation of defined neuronal networks (Singhvi et al., 1994; Branch et al., 1998, 2000; James et al., 1998; Scholl et al., 2000; Kam et al., 2001; Yeung et al., 2001).

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16.5 SEM image of a grid-shaped PDMS stamp for microcontact printing, kindly provided by Tanja Decker.

Similar to photochemical patterning methods, the procedure starts with a photolithography step to produce the mould (master stamp). Photoresists with a high aspect ratio and a thickness of more than 5 m are used to realize high relief topographical structures in the photoresist after illumination and development steps. Polydimethylsiloxane (PDMS), a viscous liquid that polymerizes to an elastomeric, rubber-like material, is poured into these molds and allowed to cure, thus solidifying to a stamp containing the topographical pattern complementary to that of the mold (Fig. 16.5). The PDMS stamp is then wet (`inked`) with a solution of organic molecules, dried and placed in contact with a surface. Similarly as with any macroscopic stamp used in everyday life, the organic molecules are transferred only at those regions where the stamp contacts the surface, leaving the pattern defined by the stamp (Fig. 16.6). As the procedure involves a drying step which may cause the denaturation of some proteins, an indirect two-step process was developed by the group of Gary Banker (Oliva et al., 2003). Bacterial protein A, which is known to have a relatively robust tertiary structure, was printed onto a tissue culture dish. Subsequently, the protein of choice ± the cell adhesion molecule L1 ± was allowed to bind to the surface from solution. Selective deposition on the pattern was achieved by coupling L1 to the Fc domain of an antibody molecule, which is known to bind protein A with high affinity. However, this two-step technique is not always necessary, as a range of matrix proteins has been shown to retain their native conformation during printing as indicated by their sustained ability to mediate cell attachment or synapse formation (Cornish et al., 2002). This fact is thought to be attributable to the formation of densely packed protein crystals during the slow drying step, which allows for the preservation of secondary and tertiary structures. The patterned surface can be in the range of several cm2 in size, while features can have an edge resolution in the sub-m range.

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16.6 Surface modification by microcontact printing. (a) Schematic representation of the steps involved: the stamp is wet with the inking solution of choice (1) and dried, leaving a thin layer of material on the stamp surface (2). Usually, a cell attractant material is used. The stamp is then pressed to the substrate surface (3), leaving the inking material only on the regions defined by the stamp topography (4). (b) SEM image of a surface onto which a grid pattern was printed using the stamp shown in Fig. 16.5. A blend of extracellular matrix protein was used as an inking solution and printed onto a polystyrene substrate.

16.4 Synaptic connections in patterned neuronal networks: communication along predefined pathways One of the most striking properties of the central nervous system lies in its ability to process and store information. Neurons communicate through electrical signals (action potentials) which originate at the cell body as a result

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of integrated signal input, and can be passed on to other cells through synaptic contacts. Synapses can be divided into two major classes, electrical and chemical synapses, which are distinct in structure and functionality. Electrical synapses, also called gap junctions, consist of a pore spanning the lipid bilayers of two adjacent cells. This pore has an internal diameter of about 2 nm, and is thus large enough for many ions and small metabolites to pass. Consequently, electrical currents can flow between the two connected cells, such that differences in the membrane potential of one cell have a direct effect on the potential of a coupled cell(s). Through this immediate electrical connection, signals are transduced directly with almost no time delay. Therefore, electrical synapses play a central role in the transmission of signals important for actions that need to be rapid, like flight reflexes. Additionally, electrical connectivity has been implicated in the synchronization of oscillatory activity in the brain (Galarreta and Hestrin, 1998). Chemical synapses are the second major type, and typically form between the axon of the presynaptic cell and a dendrite of a postsynaptic cell. They do not allow for the direct flow of current as the pre- and postsynapse are separated by a gap of about 30 nm, the synaptic cleft. If an action potential traveling down the axon of the presynaptic cell reaches the presynaptic terminal, transmitter molecules are released into the synaptic cleft. These can bind to receptors in the postsynaptic membrane, which directly or indirectly leads to the opening of an ion channel. The resulting ion flux alters the transmembrane potential and facilitates or suppresses the generation of action potentials in the postsynaptic cell (Nicholls et al., 1992). Signal transmission at a chemical synapse is unidirectional from pre- to postsynapse and occurs with a time delay of several ms. This delay arises due to the time it takes for the release of the synaptic vesicles from their anchorage in the cytoskeleton and their fusion with the presynaptic membrane. While signal transmission through electrical synapses is thought to be relatively constant, chemical synapses are subject to a wide range of modulatory mechanisms that may strengthen or weaken signal transmission in response to previous input patterns. Such activity-dependent modifications of synaptic efficacy are central on a number of capabilities of the CNS: While long-term changes are believed to be involved in learning and the formation of memory (Bliss and Lomo, 1973; Fitzsimonds et al., 1997; Ganguly et al., 2000), shortterm plasticity has been implicated in the processing of incoming signals, e.g., recognition of different input patterns or contrast adaptation (Chance et al., 1998; Kaplan et al., 2003). Sensitivity to particular patterns of sensory input is additionally attributed to features of the microcircuitry through which neurons, within a tissue, are interconnected. Different afferent signals may be attenuated or amplified depending on the `wiring' of the network, e.g., the extent to which feedback loops or reciprocal connections prevail (Muller et al., 1997; Misgeld et al., 1998; Nelson and Turrigiano, 1998; Manor et al., 1999).

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The investigation of synaptic plasticity in vitro is mostly performed by patchclamp recordings between two synaptically connected neurons in brain slices. However, the extreme complexity of neuronal connectivity within this organ renders the investigation of single contacts extremely difficult as the synaptic input from outside a constellation of interest interferes with experimentally applied signals and impedes reproducibility. Low-density neuronal cultures, in which isolated islands of two or three synaptically connected cells are found, have yielded valuable insights in synaptic modulation at the level of single synaptic contacts (Bi and Poo, 1999; Ganguly et al., 2000). Such cultures, however, do not allow for the experimental determination or alteration of the connectivity pattern, such that the impact of network architecture on signal transduction and processing cannot be studied systematically. Addressing these issues, the approach of patterned neuronal cell culture emerged. In such cultures, it is possible to grow neuronal networks of extremely low complexity, because only a limited number of pathways between the adherent cells are open for the formation of synaptic connections. Moreover, the approach also allows experimental design of the network architecture, such that linear arrangements, feedback circuits or branching pathways can be studied (Fig. 16.7). One problem with the application of patterned neuronal cultures as a model

16.7 Different possible network geometries that may be realized by growth on patterned substrates to study the impact of network architecture on signal processing. Adapted from Steward, Nature 427, 601±604 (2004).

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system lies in the fact that for a long time cells were found either to comply well with the respective surface geometry and show physiological deficits (Lauer et al., 2002; Matsuzawa et al., 2000; Stenger et al., 1998) or to develop normally while overgrowing the pattern boundaries (Ma et al., 1998; Liu et al., 2000). In particular, the formation and maturation of chemical synapses in culture fully compliant with a micropattern proved to be a challenge (Ravenscroft et al., 1998; OffenhaÈusser et al., 1997). Potential reasons for these problems may lie in the synthetic groups frequently used for surface modification which do not mediate attachment through cell adhesion molecules, but instead through other forces, e.g., electrostatic interactions. Alternatively, some of the chemical groups may have mild cytotoxic effects which impair cellular development. It has also been speculated that the geometry itself may be a factor modulating neuronal development, and that limitation of neurite outgrowth to a four-way-grid pattern is a state impeding neuronal maturation (Ravenscroft et al., 1998). Consistent with this hypothesis, Wyart and collegues reported that patterns of polylysine severely impaired cellular physiology if the adhesion sites offered for neuronal attachment were smaller than 80 m. This is an area significantly larger than a typical cell body and consequently allows the extensive outgrowth and branching of neuritis (Wyart et al., 2002). However, it was recently shown that cells highly compliant to grid patterns displayed normal physiological characteristics similar to cells grown on homogeneous control substrates. Moreover, cells were found to connect through electrical and chemical synapses that developed normally by all the criteria tested, such that communication along the predefined pattern could be observed. The pattern had been created by applying a blend of extracellular matrix proteins to a background of polystyrene by microcontact printing, such that a relatively natural substrate was offered for neurite outgrowth (Vogt et al., 2003). Moreover, the encountered synapses exhibited several forms of short-term plasticity similar to that encountered in the intact brain (Vogt et al., 2005). Short-term plasticity, the modification of synaptic strength for seconds to minutes, in response to different stimulus patterns, has been shown to be essential for a number of abilities of the mammalian brain such as the recognition of different input patterns. Encountering these features in the system while cell attachment and connectivity are highly restricted to an underlying pattern qualifies it as a suitable model system: Features basic to neuronal signal processing are preserved while the network geometry can be experimentally varied, allowing the investigation of the impact of different types of circuitry on the way incoming signals influence the basal network activity.

16.5 Conclusion Cellular adhesion to a surface can be experimentally manipulated through either chemical or topographical signals. The use of patterned surfaces consisting of

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small adhesive areas against a cell-repellent background allows the precise positioning of cell bodies thereon. In neuronal cultures, the outgrowth of neurites and thus the connectivity pattern of a neuronal network can additionally be guided by adhesive pathways interconnecting these adhesion sites. Such networks of defined geometry provide a basis for a number of biotechnological applications as well as highly defined experiments on cellular interactions, network behavior and neurocomputing. First steps towards functional networks of highly defined geometry have been taken. Highly compliant neuronal networks can be realized on patterned surfaces in which communication through mature synaptic contacts can be observed along the pattern pathways. A further challenge will be the demonstration of long-term modifications of synaptic strength in patterned networks, which are thought to underlie processes like learning and memory formation. A system reproducing these abilities of the central nervous system while network geometry connectivity pathways are experimentally controlable will be a valuable tool for the investigation of neuronal signal processing.

16.6 References Bi, G, Poo, M, Nature 401, (1999), 792±6. Bliss, T, Lomo, T, Journal of Physiology and Behavior 232, (1973), 331±356. Bohanon, T, Elender, G, Knoll, W, Koberle, P, Lee, JS, OffenhaÈusser, A, Ringsdorf, H, Sackmann, E, Simon, J, Tovar, G, Winnik, FM, J Biomater Sci Polym Ed. 8, (1996), 19±39. Branch, DW, Corey, JM, Weyhenmeyer, JA, Brewer, GJ, Wheeler, BC, Medical & Biological Engineering & Computing 36, (1998), 135±41. Branch, DW, Wheeler, BC, Brewer, GJ, Leckband, DE, IEEE Transactions on Biomedical Engineering 47, (2000), 290±300. Britland, S, Morgan, H, Wojiak-Stodart, B, Riehle, M, Curtis, A, Wilkinson, C, Experimental Cell Research 228, (1996), 313±25. Bullett, N, Short, RD, O'Leary, T, Beck, AJ, Douglas, CWI, Cambray-Deakin, M, Fletcher, IW, Roberts, A, Blomfield, C, Surface and Interface Analysis 31, (2001), 1074±76. Carter, S, Nature 208, (1965), 1183±7. Chance, F, Nelson, SB, Abbott, LF, J. Neurosci. 18, (1998), 4785±99. Clark, P, Connolly, P, Curtis, AS, Dow, JA, Wilkinson, CD, Development 108, (1990), 635±44. Clark, P, Britland, S, Connolly, P, J Cell Sci 105 (Pt 1), 203±12. Corey, JM, Wheeler, BC, Brewer, GJ, Journal of Neuroscience Research 30, (1991), 300±7. Cornish, T, Branch, DW, Wheeler, BC, Campanelli, JT, Molecular and Cellular Neurosciences 20, (2002), 140±53. Curtis, A, Wilkinson, C, Biomaterials 18, (1997), 1573±83. Curtis, A, Wilkinson, CD, J Biomater Sci Polym Ed. 9, (1998), 1313±29. Dulcey, C, Georger, JH Jr, Krauthamer, V, Stenger, DA, Fare, TL, Calvert, JM, Science 252, (1991), 551±4.

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Fitzsimonds, RM, Song, HJ and Poo, MM, Nature 388, (1997), 439±48. Fromherz, P, Schaden, H, Vetter, T, Neuroscience Letters 129, (1991), 77±80. Fromherz, P, Schaden, H, European Journal of Neuroscience 6, (1994), 1500±4. Galarreta, M, Hestrin, S, Nature Neuroscience 1, (1998), 587±94. Ganguly, K, Kiss, L, Poo, M, Nat. Neurosci. 3, (2000), 1018±26. Graf, J, Iwamoto, Y, Sasaki, M, Martin, GR, Kleinman, HK, Robey, FA, Yamada, Y, Cell 48, (1987), 989±96. Hammarback, J, Palm, SL, Furcht, LT, Letourneau, PC, Journal of Neuroscience 13, (1985), 213±20. James, C, Davis, RC, Kam, L, Craighead, HG, Isaacson, M, Turner, JN, Shain, W, Langmuir 14, (1998), 741±4. Jung, D, Kapur, R, Adams, T, Giuliano, KA, Mrksich, M, Craighead, HG, Taylor, DL, Critical Reviews in Biotechnology 21, (2001), 111±54. Kam, L, Shain, W, Turner, JN, Bizios, R, Biomaterials 22, (2001), 1049±54. Kaplan, M, Wilcox, KS, Dichter, MA, Synapse 50, (2003), 41±52. Kleinfeld, D, Kahler, K, Hockberger, P, Journal of Neuroscience 8, (1988), 4098±120. Kleinman, H, Luckenbill-Edds, L, Cannon, FW, Sephel, GC, Analytical Biochemistry 166, (1987), 1±13. Kleinman, H, Sephel, GC, Tashiro, K, Weeks, BS, Burrous, BA, Adler, SH, Yamada, Y, Martin, GR, ANN. N.Y. Acad. Sci. 580, (1990), 302±10. Kumar, A, Whitesides, GM, Applied Physics Letters 63, (1993), 2002±4. Lauer, L, Vogt, A, Yeung, C, Knoll, W, Offenhausser, A, Biomaterials 23, (2002), 3123±30. Letourneau, P, Developmental Biology 44, (1975), 92±101. Liu, Q, Coulombe, M, Dumm, J, Shaffer, K, Schaffner, A, Barker, J, Pancrazio, J, Stenger, D, Ma, W, Developmental Brain Research (2000), 223±31. Ma, W, Liu, QY, Jung, D, Manos, P, Pancrazio, JJ, Schaffner, AE, Barker, JL, Stenger, DA, Brain Research. Developmental Brain Research 111, (1998), 231±43. Manor, Y, Nadim, F, Epstein, S, Ritt, J, Marder, E, Kopell, N, Journal of Neuroscience 19, (1999), 2765±79. Matsuzawa, M, Liesi, W, Knoll, W, Journal of Neuroscience Methods 69, (1996), 189±96. Matsuzawa, M, Tabata, T, Knoll, W, Kano, M, Eur J Neurosci 12, (2000), 903±10. Merz, M, Fromherz, P, Adv. Mater. 2, (2002), 141±4. Misgeld, U, Zeilhofer, HU, Swandulla, D, Cellular and Molecular Neurobiology 18, (1998), 29±43. Mitchell, S, Emmison, N, Shard, AG, Surface and Interface Analysis 33, (2002), 742±7. Muller, T, Swandulla, D, Zeilhofer, HU, Journal of Neurophysiology 77, (1997), 3218±25. Nelson, S, Turrigiano, GG, Nature Neuroscience 1, (1998), 539±41. Nicholls, J, Martin, AR, Wallace, BG, From Neuron to Brain S Associates, Sunderland, Mass. (1992). Nuzzo, R, Allara, DL, Journal of the American Chemical Society 105, (1983), 4481. OffenhaÈusser, A, Sprossler, C, Matsuzawa, M, Knoll, W, Neuroscience Letters 223, (1997), 9±12. Oliva, AJ, James, CD, Kingman, CE, Craighead, HG, Banker, GA, Neurochemical Research 28, (2003), 1639±48. Prucker, O, RuÈhe, J, Materials Research Society Symposium Proceedings 304, (1993), 1675. Prucker, O, Schimmel, M, Tovar, G, Knoll, W, RuÈhe, J, Advanced Materials 10, (1998), 1073±7.

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Ranieri, J, Bellamkonda, R, Bekos, EJ, Gardella, JA Jr, Mathieu, HJ, Ruiz, L, Aebischer, P, Int J Dev Neurosci 12, (1994), 725±35. Ravenscroft, MS, Bateman, KE, Shaffer, KM, Schessler, HM, Jung, DR, Schneider, TW, Montgomery, CB, Custer, TL, Schaffner, AE, Liu, QY, Li, YX, Barker, JL, Hickman, JJ, Journal of the American Chemical Society 120, (1998), 12169±77. Rosenberg, M, Science 139, (1963), 411±12. RuÈhe, J, Knoll, W, Supramolecular Polymers A Ciferri, New York (2000) 565±613. Ruoslahti, E, Pierschbacher, MD, Cell 44, (1986), 517±18. Ruoslahti, E, Pierschbacher, MD, Science 238, (1987), 491±7. Sagiv, J, Journal of the American Chemical Society 102, (1980), 92. Scholl, M, Sprossler, C, Denyer, M, Krause, M, Nakajima, K, Maelicke, A, Knoll, W, OffenhaÈusser, A, Journal of Neuroscience Methods 104, (2000), 65±75. Singhvi, R, Kumar, A, Lopez, GP, Stephanopoulos, GN, Wang, DI, Whitesides, GM, Ingber, DE, Science 264, (1994), 696±8. Sorribas, H, Padeste, C, Tiefenauer, L, Biomaterials 23, (2002), 893±900. Stenger, DA, Hickman, JJ, Bateman, KE, Ravenscroft, MS, Ma, W, Pancrazio, JJ, Shaffer, K, Schaffner, AE, Cribbs, DH, Cotman, CW, Journal of Neuroscience Methods 82, (1998), 167±73. Tai, H, Buettner, HM, Biotechnology Progress 14, (1998), 364±70. Valentini, R, Vargo, TG, Gardella, JA Jr, Aebischer, P, J Biomater Sci Polym Ed 5, (1993), 13±36. Vogt, AK, Lauer, L, Knoll, W, OffenhaÈusser, A, Biotechnology Progress 19, (2003), 1562±8. Vogt, AK, Wrobel, G, Meyer, W, Knoll, W, OffenhaÈusser, A, Biomaterials 23, (2005), 3123±30. Wyart, C, Ybert, C, Bourdieu, L, Herr, C, Prinz, C, Chatenay, D, Journal of Neuroscience Methods 117, (2002), 123±31. Yeung, CK, Lauer, L, OffenhaÈusser, A, Knoll, W, Neuroscience Letters 301, (2001), 147±50.

17 Controlled cell deposition techniques C M A S O N , University College London, UK

17.1 Introduction Important challenges for the successful engineering of human tissues and organs include elucidating how the cellular microenvironment interacts with cell function ± a bidirectional, dynamic and highly complex set of interactive relationships. Thus the ability to be able to position, manipulate, orientate the architecture and control multiple cell types together with their extracellular matrix at the cellular and molecular level of resolution is paramount for the fabrication of true-to-life three-dimensional living constructs. This chapter focuses on three key areas: 1. 2. 3.

In-vivo and in-vitro cell interactions. Two-dimensional in-vitro controlled cell deposition. Three-dimensional in-vitro controlled cell deposition.

A multidisciplinary approach to the topic is mandatory since only by understanding the in-vivo embryology, histology, biochemistry and physiology together with a knowledge of cell-substrate interaction, is it possible to begin to evolve the controlled cell deposition techniques required to fabricate accurately representative two-dimensional and three-dimensional de novo cellular structures. Two-dimension controlled cell deposition has its origins in traditional cell culture where cell monolayers are seeded onto flat surfaces in sterile flasks. The history of animal cell culture dates back approximately one hundred years to the pioneering work by Harrison and Carrel (Harrison, 1907; Carrel, 1912). Over the past few decades, a paradigm shift has occurred due to the convergence of cell culture with electronic semiconductor technology allowing cells to be accurately patterned onto substrates. Three-dimensional controlled cell deposition techniques are altogether much more recent and as such are at an extremely early stage of their development. However, the advancing fields of regenerative medicine and, in particular, tissue engineering are highlighting the potential demand for such technology.

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17.2 In-vivo and in-vitro cell interactions The complex mechanisms whereby cells interact and function as tissues and organs has evolved over hundreds of millions of years. Understanding these processes in vivo in conjunction with how the same processes interact with synthetic polymers is key to controlling cell deposition in vitro.

17.2.1 Animal cells and tissues The majority of cells in the body are not isolated individual entities but are grouped together to form tissue and organs. Histologists classically classified most cells into two discrete groups; epithelial and mesenchymal cells depending upon their location and orientation. Epithelial cells principally cover surfaces (e.g. skin epithelium) and line cavities and hollow tubes (e.g. endothelium lining blood vessels). A feature common to all epithelial cells is that they form continuous sheets or layers with their outer cell membranes tightly butted up against one another like paving slabs forming a well-made pavement. In stark contrast, the mesenchymal cells (or connective tissue cells) which make up the rest of the body are widely spaced apart in a ground substance with the generic name extracellular matrix (ECM). Whilst cells are the essential unit of living material, it is the ECM which is largely responsible for giving individual tissues and organs their unique and specific properties. Thus in vivo, cells interact both with one another and with their surrounding ECM. It is therefore important that cells grown in vitro have the same orientation, density and interactions with their environment if they are to be truly representative of their in-vivo counterparts. Cells have traditionally been cultured as monolayers of single cell types which in general are not totally representative of the same cells living in a three-dimensional tissue or organ. This arrangement, however, has the advantage of simplicity and can `give important information on the characters acquired by tissues liberated from the control of the organism from which they were derived' (Carrel, 1912), e.g., physiological processes including signal transduction and the regulation of gene expression. The alternative to single cell type monolayers is to fabricate whole tissues and organs in vitro, in which collections of different cell types with a temporary artificial ECM (scaffold) are cultured together in a single bioreactor. This is the approach adopted by today's tissue engineers (Stock and Vacanti, 2001).

17.2.2 Extracellular matrix The ECM is the non-cellular structural material that surrounds mesenchymal cells in order to form tissues and organs. There are two distinct varieties of ECM. The first variety fully encapsulates mesenchymal cells, the second type is

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in the form of a thin amorphous lamina which creates the interface between the basal surface of epithelial cells and their directly underlying mesenchymal tissue (Bard, 1990). Produced by the mesenchymal cells, the ECM has a number of roles including structural support and provision of environmental signals to control site-specific cellular regulation (Junqueira and Carniero, 2003). CellECM interactions are bidirectional and dynamic in nature. Thus, mesenchymal cells directly influence the composition of their surrounding matrix whilst the matrix feeds back to the encapsulated cells. This feedback can be either biochemical or biomechanical in nature. Thus a basic understanding of cellECM interaction is vital if researchers are to be able to accurately mimic in-vivo situations in vitro. In addition, in cell culture, there is an added complexity, namely, cell-synthetic polymer interaction. The majority of animal cells are only capable of proliferating when attached to a surface, aptly named anchorage dependent cells. Therefore, in-vitro studies typically involve using a polymer surface (e.g. polystyrene) upon which the cells can anchor. To aid cell attachment, manufacturers of cell culture plastic-ware pre-treat the surfaces, since only the uppermost layers of the polymer are in direct physicochemical contact with the biological environment (Voger, 1993). In general, the adhesion of animal cells to polymer surfaces occurs in two distinct steps. Firstly, the cell on approaching a surface expands out a temporary pseudopodial extension in order to examine the surface for suitable protein ligands prior to forming an initial temporary focal attachment. This is followed by the cell spreading itself out over the surface and creating more permanent local attachments (Mosher, 1993). The adhesion process is created by a combination of non-specific and specific interactions (Voger, 1993). The nonspecific interactions include hydrogen bonding and van der Waal's forces (Parsegian, 1981), whilst specific adhesion involves receptor-ligand bonding, e.g., integrins, a family of proteins that bind to ECM molecules such as fibronectin, collagen and laminin (Irvine et al., 2002).

17.2.3 Solid-fluid interface As a generalisation, due to their tertiary structure, mammalian proteins prefer to be in an aqueous environment. However, when an aqueous solution of proteins is exposed to a solid surface, the proteins have a tendency to spontaneously accumulate (adsorb) at the solid-fluid interface due to a combination of their degree of hydrophobicity, electrical charge and polar forces. This propensity to accumulate following collision with the interface is dependent upon a number of variables, including the nature of the solid surface, temperature, aqueous solution characteristics (e.g. protein concentration and flow) and the duration of exposure (Norde, 1995; Norde and Haynes, 1996). The duration of exposure required to trigger adsorption to a synthetic surface is in the order of seconds (Vroman, 1987). Adsorption of proteins to a synthetic surface initiates the

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formation of a biopolymer film which in turn prepares the surface for cell attachment. Cell attachment quickly ensues in a matter of a few minutes. Adsorption is, however, not a simple process whereby once the synthetic surface is coated with protein the priming of the surface is totally completed. Instead, it is a dynamic process e.g. adsorbed proteins are replaced by other proteins in the solution and/or adsorbed proteins undergo conformational changes. For example, in vivo, adhesion of proteins to artificial surfaces is dominated by the Vroman effect in which the more abundant proteins are adsorbed first and then are replaced by less abundant proteins which have a higher affinity for the particular artificial surface (Vroman, 1974; Vroman and Adams, 1986). For example, when plasma is exposed to a glass surface, albumin (the most abundant protein) is adsorbed first, which is then replaced by fibrinogen (less abundant) which is in turn replaced by kininogen (least prevalent protein) (Pfeiffer et al., 1998; Jung et al., 2003). Finally, following adsorption, protein-adsorption kinetics may favour either conformational change and/or denaturing of the protein (Brash and Horbett, 1995). Understanding the cell-substrate interaction lies at the core of in-vitro controlled cell deposition, since cells in vivo are not simply stuck together with ECM `glue' in a random arrangement but are organised into diverse and distinctive patterns (Gumbiner, 1996). A detailed review of this topic has been undertaken by Saltzman (2000).

17.3 Two-dimensional controlled cell deposition techniques 17.3.1 Photolithography Photolithography creates a topographic pattern on a surface by lithography (the process of printing from a surface on which the printing areas are not elevated but are ink receptive) combined with an etching process. This approach has its roots firmly embedded in the semiconductor industry (Britland et al., 1992). Typically, the technique involves a suitable surface (e.g. borosilicate glass, fused quartz or polished silicon wafer) being spun coated with a thin photosensitive organic resist film (e.g. diazo-naphtho-quinone with a novolak base resin) (Nicolau et al., 1999). Using an exposure aligner, an opaque mask (typically a chrome film) of a desired final pattern is placed over the photosensitive resist film. Ultraviolet radiation is then directed at the unmasked areas resulting in the exposed areas of the film increasing in solubility upon the addition of a developer solution. Thus selective dissolution (wet etching) of the membrane can be achieved resulting in a photoresist pattern of membrane identical to the mask. Hydrofluoric acid is now deployed, since it can react only with the areas of exposed glass, a precise pattern is etched in the surface. The depth of the channel produced is proportional to duration of exposure. Etching of glass is isotropic, i.e., etching proceeds in all directions resulting in the undercutting of

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the mask. This results in the etched channel being `canoe' shaped in crosssection. The approximate channel width is, therefore, the width of the mask plus twice the etch depth. In contrast, if silicon is wet etched, the process is anisotropic due to the different reactivity of the different crystal planes. Overall the entire process of fabricating a chip using wet etching involves approximately 100 distinct steps (Greenwood, 2004). Whilst wet etching is the predominant etching technique there are, however, a host of alternatives, including laser ablation, micromachining, powder blasting, embossing, LIGA (Lithografie, Galvanoformung, Abformung ± lithography, electroplating, injection moulding) and plasma ion etching. The choice of etching technique depends upon a range of factors including the material to be etched, the desired channel shape, required pattern resolution and cost. Each of the methods has its advantages and disadvantages. For example, laser ablation vaporises the material and hence no undercutting of the mask, resulting in extremely precise square channels. The trade-off is that the resulting channel surface has a rough texture. Likewise, powder blasting, an anisotropic process, also results in a rough surface. Micromachining is performed by using a computer controlled hard milling wheel to erode the channels, unfortunately the size of the milling wheel limits the degree of resolution (25 microns) (Folch et al., 1999). However, achieving and maintaining conformal contact is not a trivial matter. Secondly, the surface to be patterned need not be planar due to the flexibility of the PDMS

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form. Finally, it is worth noting that the microchannels have a very high surface:volume ratio which can be both beneficial or detrimental depending upon the particular application. By flowing cell adhesion proteins or ligands through the microchannels these components are deposited resulting in a patterned surface favourable for subsequent cell attachment. A good example of the technique is by Folch and Toner (1998) who created a network of deep channels by replica moulding PDMS using microfabricated glass master moulds. Microcapillaries were created by the self-sealing action of the PDMS microstructure against the surface of a chosen substrate. Submillilitre volumes of protein were injected into these microcapillaries. The protein adsorbed only onto areas of the substrate which were exposed to the microflow. The capillaries were flushed and the PDMS microstructure carefully peeled away. Both rat collagen and human plasma fibrinogen were deployed. For cell seeding, hepatocytes were chosen since they were known not to be able to attach to the bare substrate but would readily attach to either collagen or fibronectin. Thus selective attachment of hepatocytes was achieved. Interesting, after 24 hours the hepatocytes start spreading out from the patterned areas. This was attributed as possibly being due to the adsorption of endogenously secreted ECM protein by the cells (Odenthal et al., 1992). Finally, Folch and Toner (1998), created micropatterned cocultures of hepatocytes and fibroblasts by first adding sufficient hepatocytes in solution to totally cover the protein and then adding fibroblasts in solution which readily attached to the bare substrate. As an alternative to using a protein intermediate step and if the channels are large enough, it is possible to directly flow cells through the deep microchannels and thus pattern a cell-friendly surface in a more direct manner (Folch and Toner, 1998). Folch et al. (1999) have fabricated deep (>25 micron) channels by replica moulding PDMS using microfabricated glass master moulds. Fibroblasts in solution were directly injected into the microcapillaries created by the PDMS microstructure conformally sealed against the surface of a conventional tissue culture dish. The PDMS microstructure was carefully peeled off the cell-friendly surface after two hours leaving cells arranged in the desired pattern. A microchannel network can be filled simply by placing a drop of solution at the entrance of the primary channel (service port) and capillary attraction results in the fluid being drawn into the system. However, more sophisticated hydrodynamic methods can also be used if desired including syringe drivers and peristaltic pumps (Juncker et al., 2002). The range of surfaces suitable for micropatterning using microfluidic deposition is much greater than for microcontact printing and includes gold, glass and polystyrene (Delamarche et al., 1997). Typically, microfluidics networks require submicrolitre volumes of liquids resulting in excellent economy of the reagents (Folch and Toner, 1998). This is essential as the reagents for life science microfluidic experiments are frequently extremely precious (Delarmarche et al., 1998).

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A major challenge in microfluidics systems is to create three-dimensional arrangements. To date, the complexity of fabricating such structures has prevented their construction, although this situation may be changing through the deployment of rapid prototyping technologies. For example, Chiu et al. (2000) based on the work of Anderson et al. (2000) have fabricated three-dimensional microfluidic systems and used them to pattern proteins and mammalian cells onto a surface, e.g., bovine endothelial cells. Recently, Tan and Desai (2003a,b) have also reported favourable results using microfluidic patterning to create three-dimensional biopolymer matrices of collagen-chitosan-fibronectin seeded with endothelial cells and fibroblasts. Laminar flow patterning A variant of patterning using microfluidic networks is laminar flow patterning. The principle underlying this technique is that the flow of fluids passing through the microchannels is laminar due to the small bore diameter (typically