Biomedical Surfaces (Engineering in Medicine & Biology)

Biomedical Surfaces

Engineering in Medicine & Biology Series Editors, Martin L. Yarmush and Christopher J. James Advanced Methods and Tools for ECG Data Analysis, Gari Clifford, Francisco Azuaje, and Patrick McSharry, editors Intelligent Systems Modeling and Decision Support in Bioengineering, Mahdi Mahfouf Micro and Nanoengineering of the Cell Microenvironment: Technologies and Applications, Ali Khademhosseini, Jeffrey Borenstein, Mehmet Toner, and Shuichi Takayama, editors Translational Approaches in Tissue Engineering and Regenerative Medicine, Jeremy J. Mao, Gordana Vunjak-Novakovic, Antonios G. Mikos, and Anthony Atala, editors

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Biomedical Surfaces Jeremy Ramsden

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ISBN-13: 978-1-59693-240-1

Cover design by Igor Valdman

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10 9 8 7 6 5 4 3 2 1

Contents 1 The adsorption and desorption of biomacromolecules 1.1 Elementary physico-chemical considerations . . . . . . 1.2 Monolayer adsorption . . . . . . . . . . . . . . . . . . . 1.2.1 Effect of flow rate . . . . . . . . . . . . . . . . . 1.3 Desorption . . . . . . . . . . . . . . . . . . . . . . . . . 1.4 Adsorbed protein exchange . . . . . . . . . . . . . . . . 1.5 Multilayer adsorption . . . . . . . . . . . . . . . . . . . 1.6 Multicomponent adsorption . . . . . . . . . . . . . . . 1.7 Adsorption onto heterogeneous surfaces . . . . . . . . . 1.8 Further reading . . . . . . . . . . . . . . . . . . . . . . 2 Chemical and physical attributes of biomedical 2.1 Interaction energetics . . . . . . . . . . . . . . . 2.1.1 Pairwise interactions . . . . . . . . . . . 2.1.2 Protein surface heterogeneity . . . . . . 2.1.3 Three-body interactions . . . . . . . . . 2.2 Wettability . . . . . . . . . . . . . . . . . . . . 2.2.1 The wetting transition . . . . . . . . . . 2.2.2 Wetting dynamics . . . . . . . . . . . . . 2.2.3 Dewetting dynamics . . . . . . . . . . . 2.3 Morphology . . . . . . . . . . . . . . . . . . . . 2.3.1 Roughness . . . . . . . . . . . . . . . . . 2.4 Chemistry . . . . . . . . . . . . . . . . . . . . . 2.4.1 The principles of molecular recognition . 2.4.2 Corrosion . . . . . . . . . . . . . . . . . 2.5 Biodegradability . . . . . . . . . . . . . . . . . . 2.6 Further reading . . . . . . . . . . . . . . . . . .

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surfaces . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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31 32 32 35 37 39 40 41 41 42 42 46 46 51 53 54

3 Biofluid compatibility 55 3.1 Haemocompatibility . . . . . . . . . . . . . . . . . . . . . . . 56 3.1.1 The composition of human blood . . . . . . . . . . . . 56 i

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CONTENTS

3.2 3.3 3.4 3.5 3.6 4 Cell 4.1 4.2 4.3 4.4 4.5 4.6

4.7

3.1.2 Initial recognition events . . . . . . . 3.1.3 The coagulation cascade . . . . . . . 3.1.4 The immune system . . . . . . . . . 3.1.5 The innate immune system . . . . . . 3.1.6 The adaptive immune system . . . . 3.1.7 Haemodialysis to combat renal failure 3.1.8 Stents . . . . . . . . . . . . . . . . . Urine compatibility . . . . . . . . . . . . . . Lachrymal secretions . . . . . . . . . . . . . Buccal secretions . . . . . . . . . . . . . . . Bioanalytical surfaces . . . . . . . . . . . . . Further reading . . . . . . . . . . . . . . . . and tissue interaction with surfaces Principles of eucaryotic cell architecture Cell adhesion . . . . . . . . . . . . . . . The adhesion of cells to artificial surfaces Principles of tissue architecture . . . . . 4.4.1 Inflamed tissue . . . . . . . . . . Assimilation . . . . . . . . . . . . . . . . Bacterial infection of implants . . . . . . 4.6.1 Bacterial structure and dynamics 4.6.2 Biofilms . . . . . . . . . . . . . . 4.6.3 Bacterial adhesion . . . . . . . . Further reading . . . . . . . . . . . . . .

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5 Lubrication and wear 5.1 Fundamentals of tribology . . . . . . . . . . . . . . . . . . . 5.1.1 Dry friction . . . . . . . . . . . . . . . . . . . . . . . 5.1.2 Lubrication . . . . . . . . . . . . . . . . . . . . . . . 5.2 Friction in hip endoprostheses . . . . . . . . . . . . . . . . . 5.3 Possible physiological effects of particles released as a result of wear . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.3.1 Physical properties of (nano)particles . . . . . . . . . 5.3.2 Chemical properties of (nano)particles . . . . . . . . 5.4 Further reading . . . . . . . . . . . . . . . . . . . . . . . . . 6 Surface generation, engineering 6.1 Machining . . . . . . . . . . . 6.1.1 Mechanical operations 6.1.2 Laser micromachining

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58 58 61 62 62 65 66 66 67 68 68 69

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91 93 93 95

coating 97 . . . . . . . . . . . . . . . 99 . . . . . . . . . . . . . . . 100 . . . . . . . . . . . . . . . 101

CONTENTS

6.2

6.3

6.4

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6.1.3 Electrochemical methods . . . . . . . Surface engineering . . . . . . . . . . . . . . 6.2.1 Plasma modification . . . . . . . . . 6.2.2 Ion implantation . . . . . . . . . . . 6.2.3 Anodizing . . . . . . . . . . . . . . . Surface coatings . . . . . . . . . . . . . . . . 6.3.1 Physical vapour deposition (PVD) . 6.3.2 Chemical vapour deposition (CVD) . 6.3.3 Ceramic coatings by plasma spraying 6.3.4 Self-assembled coatings . . . . . . . . Further reading . . . . . . . . . . . . . . . .

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7 The metrology of biomedical surfaces 7.1 Energetics . . . . . . . . . . . . . . . . . . . . . . . . . . 7.1.1 Wetting . . . . . . . . . . . . . . . . . . . . . . . 7.1.2 Probe particles . . . . . . . . . . . . . . . . . . . 7.2 Morphology . . . . . . . . . . . . . . . . . . . . . . . . . 7.2.1 Contact techniques . . . . . . . . . . . . . . . . . 7.2.2 Porosity . . . . . . . . . . . . . . . . . . . . . . . 7.2.3 Non-contact (optical) techniques . . . . . . . . . . 7.2.4 Layer thickness and density . . . . . . . . . . . . 7.3 Dynamical structure measurements . . . . . . . . . . . . 7.4 Chemical surface analysis . . . . . . . . . . . . . . . . . 7.4.1 Layer chemistry and multicomponent distribution 7.4.2 Electrical surface properties . . . . . . . . . . . . 7.4.3 Protein adsorption and desorption . . . . . . . . . 7.5 Cytometry . . . . . . . . . . . . . . . . . . . . . . . . . . 7.5.1 Microarrays . . . . . . . . . . . . . . . . . . . . . 7.5.2 The polymerase chain reaction . . . . . . . . . . . 7.5.3 DNA sequencing . . . . . . . . . . . . . . . . . . 7.5.4 Two dimensional gel electrophoresis (2DGE) . . . 7.5.5 Proteolysis and chromatography . . . . . . . . . . 7.6 Quantifying biocompatibility . . . . . . . . . . . . . . . . 7.6.1 Effects of roughness on response . . . . . . . . . . 7.7 Further reading . . . . . . . . . . . . . . . . . . . . . . . 8 Prospects for smart and nano materials 8.1 The nature of responsiveness . . . . . . . 8.2 Bioactive coatings . . . . . . . . . . . . . 8.2.1 Cell adhesion-promoting coatings 8.2.2 Call adhesion-inhibiting coatings

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103 104 104 105 106 106 107 108 110 111 114

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iv

CONTENTS

8.3 8.4 8.5

8.2.3 Drug delivery surfaces . . . . . . . . . . . 8.2.4 Biocidal surfaces . . . . . . . . . . . . . . Carbon nanotubes in biomedical surfaces . . . . . A complexity sciences perspective for biomaterials Further reading . . . . . . . . . . . . . . . . . . .

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152 152 153 154 154

List of Tables 1

2

Table of the relative importance (ranked by numbers of occurrences of the listed words) of body parts improved by prostheses. Source: Prostheses List, Australian Health Insurance Association (AHIA), 2005. . . . . . . . . . . . . . . . . . . . .

7

Table of the relative importance (ranked by numbers of occurrences of the listed words) of implanted devices. Source: Prostheses List, Australian Health Insurance Association (AHIA), 2005. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

8

3

Table of the relative importance (ranked by numbers of occurrences of the listed words) of the materials used in implanted devices. Source: Prostheses List, Australian Health Insurance Association (AHIA), 2005. . . . . . . . . . . . . . . . . . . . . 11

4

Table of typical applications of biomedical materials. . . . . . 12

2.1

Surface tensions/(mJ m−2 ) and Hamaker constants of of a selection of substances. Surface tension data mostly from van Oss (1996); Hamaker constants from Dahneke (1972), van Oss et al. (1988) and Roth et al. (1996). . . . . . . . . . . . . . . . 36

2.2

Table of common roughness parameters. hi is the height of the ith spot on the surface, with reference to some arbitrary fixed level; N is the number of measured spots. Additional parameters are the skew or skewedness (the third moment) and the kurtosis (the fourth moment). . . . . . . . . . . . . . 43

2.3

Truth table for the complementation operator ⊗. d denotes a hydrogen bond donor, a a hydrogen bond acceptor, and n neither. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 47

2.4

The electrochemical series of metals used in biomedical applications. Standard potentials are given versus the hydrogen electrode (standard potential = 0). . . . . . . . . . . . . . . . 52 v

vi

LIST OF TABLES 3.1

3.2 3.3 3.4

The composition of human blood. Data is for whole blood, except the protein concentrations, which are given for plasma (serum), i.e. blood from which the cells have been removed. . The principle blood cascades. . . . . . . . . . . . . . . . . . The composition of human urine. . . . . . . . . . . . . . . . The composition of human tears. . . . . . . . . . . . . . . .

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57 61 70 71

4.1

The molecules expressed during the successive stages of leucocytes homing in from the bloodstream to the vascular endothelium. Where ligand . . . . . . . . . . . . . . . . . . . . . 79

5.1

The composition of synovial fluid (major (by quantity) components). . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 90

7.1

Surface tensions/(mJ m−2 ) of a selection of liquids useful for surface tension measurements. Viscosity (η) and boiling point are given because of their practical importance in choosing liquids for the experimental measurement of contact angles. . . 119

List of Figures 1

The relationship between applications and attributes. . . . . . 16

1.1

Sketch of the interfacial potential energy acting between a compact macromolecule approaching an adsorbent surface. The abscissa gives the separation between the macromolecule and the surface, and the ordinate the potential energy. Far from the surface (and for the physiological situations relevant to this book, that may mean as little as 10–20 nm) there is no perceptible interaction, and the molecule moves with the motion of an unbiased random walk. If it happens to approach the surface, it will typically sense an increasingly repulsive force. As the molecule approaches yet closer, the repulsive force is increasingly counteracted by a shorter range attractive force (this is the quasi-universal Lifshitz-van der Waals interaction), which at very close range (a few nanometres) increasingly dominates, and those proteins that manage to diffuse over the thermal barrier (height ∆Ga ) peaking at the crossover between the long-range repulsive and short range attractive forces fall into a more or less deep potential well, whose depth ∆Gd is determined by the onset of the extremely short-range Born repulsion. . . . . . . . . . . . . . . . . . . . 19

1.2

Sketch of the adsorption process for a typical globular protein (after Fern´andez and Ramsden, 2002). . . . . . . . . . . . . . 20

1.3

Sketch of the events associated with the approach, arrival and residence of a protein to, at and on a biomedical surface. . . . 21

1.4

Convective diffusion in a blood vessel. v is the velocity of the liquid. The diffusion boundary is indicated by the dashed lines. The dotted line indicates a hypothetical layer whose thickness is of molecular dimensions. . . . . . . . . . . . . . . 26 vii

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LIST OF FIGURES 2.1

Small droplets of a liquid L sitting at equilibrium on a horizontal surface S in the presence of their vapour V. (a, left-hand side) on a relatively hydrophobic surface; (b, right-hand side) a relatively hydrophilic surface (assuming L = H2 O). . . . . . 40 2.2 Wear of the metallic head and ferulene bush of a novel endoprosthesis 10 years after initial implantation (source: Sytenko Institute of Spine and Joint Pathology, Kharkov. . . . . . . . . 53 3.1 3.2

Schematic diagram of a molecular cascade. . . . . . . . . . . . 59 Examples of modern stainless steel stents. . . . . . . . . . . . 66

4.1

Sketch of the so-called lipid bilayer membrane showing peripheral and integral membrane proteins and glycoproteins. H, lipid head-group; T, lipid tail; P, peripheral membrane protein; A, protein and ca, comprised of either an acyl chain, or hydrophobic amino acid residues, or both; E, extracellular domain of an integral membrane (glyco)protein; Tr, transmembrane portion of the integral membrane protein, typically comprising one or more of the alpha helices traversing the membrane; C, cytoplasmic domain, which may be simply a loop connecting the transmembrane portions. . . . . . . . . . 74 Sketch of a eucaryotic cell. . . . . . . . . . . . . . . . . . . . . 76 Sketch of a spread eucaryotic cell (in cross-section). On the left the overall shape, on the right detail of the cell-substrate interface. Quite a lot of detail is known about the molecular constituents of the filopodia. . . . . . . . . . . . . . . . . . . . 77

4.2 4.3

5.1 5.2

Dry friction. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 88 Lubricated friction. At very low velocities one refers to boundary lubrication, or nanofriction, because the lubricant film is less than 100 nm thick. More exactly, the ratio h/R of the thickness h of the lubricating film to the roughnesses R of the surfaces is low. Around the minimum of the curve one has mixed film (elastohydrodynamic) lubrication. At very high velocities hydrodynamic lubrication dominates: the friction is determined merely by the viscosity of the lubricant; the surfaces are no longer in contact with one another. . . . . . . . . 89 5.3 Friction coefficients of (from top to bottom and five main): αalumina ceramic/polyethylene; Co-Mo-Cr alloy/polyethylene; and monocrystal sapphire/polyethylene. . . . . . . . . . . . . 91

LIST OF FIGURES

ix

5.4

Wear of the metallic head and ferulene bush of a normal endoprostheses after 10 years of operation. . . . . . . . . . . . . 91

5.5

Electron micrograph by SEM of Co-Cr particles (0.3–5 µm) retrieved from tissues adjacent to a human implants at revision surgery. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 92

5.6

Electron micrograph by ESEM of a microparticle (centre) and nanoparticles (top left, bottom centre) in the liver of a rabbit following implantation in the femoral bone marrow. The particles used are those shown in the previous figure. . . . . . 92

5.7

Sketch of the energy levels of a semiconductor. CB, conduction band; VB, valence band; ED, electron donor level (due to an impurity or defect). . . . . . . . . . . . . . . . . . . . . . . . . 92

6.1

Tissue removal tool for use with an endoscope directly manufactured from a solid block in under 15 minutes by 5-axis machining (courtesy of Kern GmbH). . . . . . . . . . . . . . . 100

6.2

Brass mould insert for the injection moulding of microfluidic devices produced by micromilling (Gottschlich). . . . . . . . . 100

6.3

Micro cochlear implants fabricated by injection moulding, compared with the head of a safety match (courtesy of Battenfeld). 100

6.4

(a) Laser interference surface texturing on a steel mould surface; (b) the texturing replicated in polypropylene (courtesy of Groenendijk and Meijer). . . . . . . . . . . . . . . . . . . . 102

6.5

Lupin leaves, showing ultrahydrophobicity. . . . . . . . . . . . 102

6.6

Lens haptics (holders for replacement eye lenses) etched from 0.125 mm thick Kapton (polyimide) film (courtesy of TechEtch Inc., Mass.). . . . . . . . . . . . . . . . . . . . . . . . . 103

6.7

So-called “painless” injection needles machined in silicon using photolithography and chemical etching (courtesy of Georgia Institute of Technology, USA). . . . . . . . . . . . . . . . . . 103

6.8

Polished sidewall of a NiTi foil produced by electrolytic photoetching. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 103

x

LIST OF FIGURES 6.9

Scanning electron micrographs of titanium anodized in acetic acid showing two different types of surface morphology: (a) a non-porous barrier type in the control implant; and (b) a porous type in the test implant. (c) shows the honeycomb appearance of the porous implant at tenfold higher magnification (the scale bar in the top left-hand corner represents 10 µm). After Sul, Y.-T. (2002). The porous surfaces exhibit better histomorphometrical parameters (bone-implant contact and quality of newly formed bone) and osteoconductivity was more pronounced around the implants. . . . . . . . . . . . . . 106

6.10 Trabecular metal (courtesy of Zimmer). See text for more details. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 110 6.11 Graded coating. See text for more details. . . . . . . . . . . . 111 6.12 Sketch of the polyelectrolyte deposition procedure. . . . . . . . 112 6.13 The nanoscopic structure of seashells. Dense, hard inorganic particles are bonded together by bioorganic polymers. The volume fraction of the latter is typically only a few percent. It is believed that the particles and the polymer are bound by opposite electrostatic charges. . . . . . . . . . . . . . . . . . . 114 7.1

Diagram of an atomic force microscope (not to scale). The sample is fixed on the platform P, in relation to which the block B can be moved in the x, y and z directions. A flexible cantilever is mounted on the block, and on the end of the cantilever is a sharp tip. The beam from a diode laser L is directed onto the cantilever, from which it is reflected onto a split photodiode D. The block is brought up to the sample such that the tip is almost in contact with it, and then scanned across the sample, i.e. in the x, y plane. Sample declivities cause the tip to move up and down, and the corresponding deflexions of the cantilever are faithfully recorded via the ratio of the signals from the two halves of the photodiode. Sophisticated signal processing allows the three dimensional topography of the surface to be extracted. By using a more elaborate setup with a quadruple split photodiode, sideways deflections of the cantilever can also be recorded, enabling the friction coefficient of the sample surface to be quantified at the nanometer scale. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 124

Preface 7.2

7.3

7.4 7.5

7.6

8.1 8.2

xi The principles of scanning near field optical microscopy (SNOM). Left hand diagram: SNOM in illumination mode: a dielectric probe illuminates the sample from above, the light is collected in the far field; dashed line: reflexion mode in which the radiation reflected off the surface is collected by far field optics (alternatively the probe tip itself could be used). Right-hand diagram: SNOM in collection mode: the sample is illuminated from below and a dielectric probe collects the light coming through the sample. The two critical requirements are: (i) a sub-wavelength light source, and (ii) placing the sample in the near-field zone of the light source. . . . . . . . . . . . . . . . . 127 Polychromatic speckle patterns for (a), upper panel, Ra = 0.4µm; and (b), lower panel, Ra = 6.4µm. In the former case, the speckles are elongated, and in the latter, completely decorrelated. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 128 Prototypical force-displacement cycle on the atomic force microscope. See text for full explanation. . . . . . . . . . . . . . 130 Characteristics, and advantages and disadvantages, of the various optical techniques available for thin surface film investigation. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 136 Sketch of the types of contact between a cell and a solid planar substrate. a, a cell cultured in bulk medium that has just landed on the surface, and still retains its spherical shape; b, an idealized spread cell of the same volume as its spherical precursor, and making closely parallel contact but the substrate; c, a more realistic representation of a spread cell, showing the somewhat irregular shape of the cytoplasm, and the presence of filopodia some tens of nanometres long between the main body of the cell and the substrate. The DNA-rich nucleus of the cell is represented by the hatched zone. . . . . . . . . . . . 139 Diagram showing information flows in a general responsive system. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 149 Sketch of polyelectrolyte chains tethered by their ends to a substrate (hatched). (a) collapsed state; (b) extended state induced by adding a strong acid HA, which ionizes and protonates the amine groups, causing them to become electrostatically charged. . . . . . . . . . . . . . . . . . . . . . . . . . . . 150

xii

Preface

Preface The subject of biomaterials—from which biomedical surfaces are made— has grown up in a very ad hoc manner. There is no universal agreement about the definition of biomaterials. Some authors restrict the term to describing the materials from which prostheses and other medical implants are made—i.e. materials with distinctively medical applications. Others includes materials for drug delivery, biosensors, protein separations etc.—anything indeed which comes into contact with biomolecules, hence also including ex vivo applications for analysis and production. If this definition is extended to biomatter in general, then even textiles on against the skin would qualify to be called biomaterials. Yet others include the living matter itself in the definition—even biomolecules like DNA. Adding to the confusion is the fact that most biomaterials in use at present have been adapted from other, non-medical and non-biological, purposes. For example, titanium was initially developed for rocket motors, silicone rubbers for industrial adhesives, polytetrafluoroethylene textiles for heavy-duty water repellent clothing. All these novel materials were subsequently adapted for medical use. During this adaptation a great deal has of course been learnt about interfacing synthetic materials with living tissue. The time is now ripe for exploiting this knowledge in combination with the unprecedented control over the constitution of materials now being enabled by nanotechnology. Hence, for the first time, the fully rational design of biomaterials becomes a possibility. This book is conceived to usher in this new age of rational biomaterials design. The surface of any component in a device destined for in vivo implantation or for being placed in contact ex vivo with biomolecules (note that some applications, such as the tubing of haemodialysers, fall between these two categories) plays a crucial role in its performance. It is through its surface that the component interacts with the biological world with which it is in contact. This book focuses on those surfaces. One of the core premisses is that whenever an artificial material is brought into contact with living tissue or fluid containing dissolved proteins, adsorption of proteins on the surface of the material is the first significant event to occur. Deep understanding of this adsorption phenomenon is therefore essential. Therefore, after the Introduction, the first numbered chapter deals with the fundamentals of the interaction of biomacromolecules with surfaces. This knowledge will be used directly in later chapters dealing successively with biofluids in contact with materials, and with the interaction of living cells and tissues with surfaces. The chapter immediately following that on

Preface

xiii

biomacromolecular adsorption is also of a fundamental nature, dealing with the attributes of biomedical surfaces from a somewhat abstract viewpoint. Lubrication and wear, often leading to the release of immunogenic particles into the body, are discussed in a separate chapter. The processes of manufacturing biomedical devices places certain constraints on their surface finishes, and an extensive body of expertise has been built up in order to find suitable compromises between the demands of fabrication and the demands of biocompatibility; this knowledge is summarized in Chapter 6. The metrology of surfaces is particularly important for controlling their quality during manufacture, but it is also an essential research tool for advancing fundamental understanding of how surface attributes are correlated with biological responses. Despite numerous attempts, as yet the science appears to be some considerable distance from a proper theory of these responses, and overall mastery of the field has not been helped by premature assertion of propositions whose basis is in fact rather weak. Just as that most excellent of biomaterials, bone, is fashioned through the concerted action of the bone-constructing osteoblasts and the bone-destroying osteoclasts, so should we also remember that the biomaterials field also has need of constant destruction of erroneous ideas in order to become a strong edifice. The final chapter of the book is oriented towards the future and includes a discussion of promising directions for smart materials of various kinds, such as bioactive and biodegradable surfaces. Nanotechnology applied to biomaterials is a particularly important development for biomedical materials, since the possibilities of achieving unusual combinations of material attributes are greatly enhanced. This should enable a more perfect practical realization of desirable design principles than has hitherto been achievable. It is already well appreciated that many disciplines are needed for work in biomedical surfaces. Materials science and engineering are obviously essential, but a good knowledge of biology is also required. Almost everything in biomaterials depends on the interface between the living and the nonliving—the bio/abio interface, which may simply be called the problem of biocompatibility, and which in turn fundamentally depends on the biomolecular interactions dealt with in Chapter 1. A complicating feature of biomaterials is the fact that the living matter is constantly modifying the non-living, as well as itself, and useful insight can be gained by considering the process as being driven by an active flow of information between the living and the non-living components. This aspect is briefly touched upon in Chapter 8. There are, I suppose, two ways in which to write such a book as this. One is to start from the viewpoint of the engineer or materials scientist, who will be well acquainted with materials and how to modify them, but may have only a rudimentary knowledge of biology, even its more chemical

xiv

Preface

aspects, such as the structure and dynamics of biological macromolecules; and the other is to start from the viewpoint of the biologist, or the medical scientist, who may have a great deal of knowledge about living organisms and their molecular constituents, but whose knowledge of materials is likely to be based on a few elementary notions propagated heuristically in the literature, and not always based on solidly demonstrable principles. On the basis of the time-honoured Principle of Sufficient Reason, one may suppose that the field is divided roughly equally between these two camps, and thus it would appear to be a hopeless task to attempt to cover the ground in such a way that both sides will be equally engaged. I have set out to do so by maintaining an abstract overview as far as his possible, focusing on fundamental aspects and underlying concepts that may be unfamiliar to both sides. For detailed, specific facts the reader is invited to consult some of the encycopaedic texts available elsewhere, such as the monumental work “Biomaterials” edited by Buddy Ratner and colleagues, or the enormous research and review literature, that is too vast to be reviewed other than superficially in a book of this size. The present book is therefore not a “how to” manual; in contrast, it seeks to uncover the general principles that lie beneath the vast web of facts, which should thereafter be understood in more compact form; in other words, it provides a framework with which the already vast accumulation of facts can be made more intelligible. It is also hoped that fruitful lines of future research will become apparent. To reiterate, a very important guiding principle developed in this book is that all materials are modified by protein deposition once they are inserted into a living body. Much of the science and technology of biomaterials has been developed simply for dealing with that. Another pervasive concept is that the environment of a biomaterial is very much a system, in the sense of being constituted from inseparable interacting components that must be considered in their entirety in order to be able to formulate a proper description of the body’s response to the biomaterial, and the biomaterial’s response to the body. This book could well have been subtitled “the principles of biocompatibility”, but biocompatibility itself has no unique, consensus definition in the literature. The most compact definition of biocompatible is ‘tolerant of life’, but one very often finds definitions (of biocompatible) such as “the ability of a material to perform with an appropriate host response in a specific application” (from the Williams Dictionary of Biomaterials). Now, as Sommerhoff has so clearly explained, “the notion of adaptation when applied to living nature refers to the widespread and striking appropriateness which organic activities show in relation to the needs of the organism, and to the effectiveness with which organisms meet the demands made upon them by their

Preface

xv

environment.” Thus adaptedness implies appropriateness, but goes beyond it because the response must not only be appropriate for the particular set of concrete circumstances with which the practitioner is inevitably engaged in any particular instance, but also appropriate had the initial circumstances been somewhat different. Given the variability inherent in the living environment, this is an extremely important point. For example, a stent might be designed to perform very well under a particular set of environmental variables including temperature, chemical composition of the blood, and flow rate. Very often, indeed, stent performance is optimized in the laboratory under a very narrowly defined set of conditions. Clearly, if the stent malperforms whenever one or more of these variables are changed, such as is very likely to happen during the life of a patient, the stent cannot be said to be adapted. This is a particular challenge for the designer of biomedical surfaces, and which may well exceed the challenges to engineers in other fields, in which environmental variation tends to be more constrained and predictable. A note about referencing. Under the “Further reading” sections that close each chapter will be found a list of books and review articles that have either been consulted during the preparation of the chapter, or contain useful additional material for the reader. References to the primary literature are generally restricted to important recent research papers that are not yet discussed elsewhere, or papers that can be recommended in order to fill a gap in coverage, inclusion of which the available space does not permit. In this age of highly effective internet search engines and specialized electronic databases, the reader should have no difficulty in gathering more information on most of the topics brought up in the book. I owe a great debt of gratitude to my colleagues at Cranfield—David Allen and David Stephenson for discussions about the manufacture of surfaces, Jeff Alcock for discussions about lubrication—and Gert Goch and Stephan Patzelt from the Universit¨at Bremen for having called my attention to photoacoustic microscopy as a useful method for surface characterization, and for having explained speckle to me, respectively.

xvi

Preface

Introduction

1

Introduction The increasing trend to incorporate artificial devices into the human body, or otherwise intimately associate them with the body as in haemodialysis, has sharply focused attention onto the compatibility of the materials from which such devices are made with human physiology. Given that it is widely predicted that more and more functions will be taken over by artificial devices, not only for repairing damaged function but also for enhancing function, the need to solve the problem of how to design and manufacture such materials to make them maximally compatible with living tissue (including biofluids such as blood) has become a major priority of the medical device industry. Surfaces are the primary place of contact between a biomaterial and its host organism. A ‘biomedical surface’ is defined as the surface of the device used in the medical application, which obviously covers implants in a living organism, but also analytical devices used in clinical testing. The adsorption of biomacromolecules at the surface is pivotal for biocompatibility. The impossibility of keeping proteins away from most implants means that very careful consideration has to be given to this aspect, and both prevention and promotion for, respectively, bloodstream implants and bone replacement and repair are required. Although these two environments (bloodstream and bone) are often presented as diametrically opposed, the surgery involved in implanting prostheses into bone or soft tissue is usually invasive and an implant designed for tissue will almost inevitably first come into contact, albeit briefly, with blood, a factor that must be taken into account in its design. Furthermore, tissue cells are highly capable of synthesizing adequate quantities of extracellular matrix proteins, and the general response to the presence of an artificial surface inside the body is for the cells in its vicinity to synthesize and excrete proteins, which are adsorbed on the surface, in order to make that surface biocompatible. The dictionary definition of ‘biocompatible’ is ‘tolerant of life’. A more convenient definition for the materials scientist is ‘responding appropriately to a given physiological situation’, but although amenable to laboratory testing, this may be too restrictive for practical use: a properly useful definition is ‘adapted to its physiological environment’, i.e. able to cope with the full variety of conditions prevailing during its lifetime, or that of its host. This may include playing an active role in the host physiology, much as a pharmaceutical product (a drug). One should also bear in mind that biocompatibility

2

Introduction

should properly also include the ability to resist bacterial infection, one of the current problems of prostheses; an implant surface may therefore have to be intolerant of bacterial life. As far as application to bio-analytical devices is concerned, ‘biocompatibility’ simply needs to mean ‘not altering the composition of any analyte’, e.g. by selective adsorption, or by leaching something out of itself. At one time, it was felt that bio-inertness was the goal of any designer of a biomedical device. ‘Bio-inert’ can be defined as ‘not giving rise to any perceptible reaction by the living material’, and it was presumed that in order to achieve bio-inertness, the device should be chemically as unreactive as possible, and physically as devoid of any morphological features as possible. These attributes in turn should have led to biological inertness—non-toxic and non-immunogenic materials. The non-toxicity and non-immunogenicity provide an operational definition of bio-inertness. Materials such as polytetrafluoroethylene (PTFE) with minimal roughness would fulfil the criteria of chemical unreactivity and minimal physical (topographical) features. Nevertheless, PTFE typically does provoke a host response, because of the reason given above: it is a substrate for protein adsorption, and the adsorbed proteins, which typically undergo conformational changes, are recognized as foreign by the immune system and promptly attacked. Furthermore, while almost atomically smooth surfaces are now achievable by ultraprecision machining, such minimal roughness is expensive and needs to be justified in terms of effectiveness. Thus, bio-inertness is no longer seen as the ultimate in biocompatibility; the goal is rather active management of biocompatibility through understanding the bio/abio interface and anticipating the range of host responses likely to occur. In a nutshell, this book is about how to achieve that management. Prostheses have to fulfil demanding structural and mechanical requirements, and the material best for those functions may be bio-incompatible. Surface treatment or coating provides a means to overcome that problem, which means both integration within the host physiology and stabilization with respect to corrosion and wear. As already emphasized, in nearly every case, it is the surface of the artificial object that constitutes its interface with the living tissue. Therefore, it is appropriate to devote some attention to the practical means of modifying the surfaces of the materials from which devices are made (Chapter 6). Practitioners in the field sometimes assert that biocompatibility is nonquantifiable. It is hard to justify this point of view, according to which the field ceases to be a science in the generally accepted sense. To be sure, it is very difficult to quantify biocompatibility. Two of the contributors to Ratner et al.’s “Biomaterials” laconically remark that “The regulation of cell

Introduction

3

function by substrate surface microtexture may be a complex affair.”1 It has long been a goal of biomaterials research to find correlations between materials properties and host response. The results from the many investigations undertaken with this goal in view have been rather mixed. For example, it is often been proposed that the hydrophobicity of a surface, or its converse, hydrophilicity, determine biocompatibility as measured by the ability of cells to spread on the surface. Taylor however showed that the spreading of conjunctiva cells in Eagle’s medium is excellent on glass and cellulose acetate, which lie at opposite ends of the spectrum of hydrophobicity, and also excellent on a layer of adsorbed fibrinogen, which has an intermediate degree of hydrophobicity. Yet a layer of adsorbed albumin, with hydrophobicity intermediate between those of glass and fibrinogen, rather inhibits spreading; it is better on Teflon, which is more hydrophobic, yet completely inhibited on the even more hydrophobic adsorbed silicone.2 There is of course a priori no reason to suppose that the relationship is monotonic, but in the absence of any underlying theoretical justification, and hence prediction except in the vaguest and most qualitative of terms, we do not really know what behaviour to expect. It is a great difficulty that both material properties and host response are multidimensional quantities very difficult to subsume into a single parameter, which is possibly the origin of the idea that quantification is unachievable. The continuous technological advances increasingly enabling the characterization of materials at the nanoscale, and the equivalent characterization of biological response at the molecular level, is however dramatically changing the situation. Therefore, it is legitimate to hope that we stand at the threshold of a real breakthrough in achieving the long sought-for correlation, which doubtless will not be simple, but at least it will be capable of quantitative prediction. Meanwhile, we should note that materials properties can already be characterized by many quantitative parameters, in some cases in well-established ways (see Chapter 7). The quantification of host response is more difficult, not least because it is continuously evolving, but modern technology allows such detailed characterization as the determination of the spectrum of expressed proteins (the transcriptome or the proteome), as well as the instantaneous concentration of metabolites (the metabolome), etc.

Areas of application 1

J.A. Jansen and A.F. von Recum, p. 222. A.C. Taylor, Attachment and spreading of cells in culture, Experimental Cell Research, Suppl. 8 (1961) 154–173. 2

4

Introduction

Implants are defined as objects in contact with living tissue, and longterm implants are defined as objects in contact with living tissue more-or-less permanently. In reality, that might only mean several years, following which a fresh operation to replace the implant (or remove it) is necessary. These repeated interventions are, however, extremely problematical, especially in the case of elderly patients, for whom each successive intervention becomes more and more life-threatening. Therefore, it is a major current challenge in biomaterials to develop implants that are truly permanent, and one of the keys to doing so as to manage the interface between the implant and the living host. Examples of this kind of application are bone and joint replacements, dental prostheses, implanted sensors (e.g. for blood glucose), stents, heart pacemakers and heart valves. Apart from direct effects on host physiology, long-term implants are frequently subject to bacterial colonization, and to wear. Bacterial infections tend to involve the formation of biofilms and as a result to become extremely persistent, the bacteria becoming resistant to all usable antibiotics. In relatively benign cases, deleterious effects may be confined to local inflammation, but in more severe cases, systemic effects may arise, and often the infection can only be halted by removing the prosthesis. Wear, especially of artificial joints, has an indirect effect on host physiology by releasing particles into the body. These particles they become widely distributed throughout the host. Their highly curved surfaces may be substrates for adsorbing and denaturing proteins, with the result that they become recognized by the immune system as foreign bodies, and an immune reaction of greater or lesser severity is initiated, often leading to problematic inflammation (possibly systemic) and even cancer. Many of these reactions are already known through the study of diseases (e.g. silicosis) caused by inhaling airborne dusts. Medium-term implants are defined as objects in contact with living tissue for a limited duration, which may be as little as a few days. The exigencies of biocompatibility are much less severe than in the case of long-term implants, simply because they are fairly rapidly removed. Examples of such implants include sutures and drug delivery particles. Small devices introduced into the body for surveillance or repair, such as the swallowable capsule-cameras for surveying the gastro-intestinal tract, also fall into this category. Ex vivo, one has tissue scaffolds, e.g. acting as supports for cells grown for skin replacement and reconstruction, and sometimes the artificially grown tissue may be introduced into the body together with its scaffold. For this type of implant, a very promising approach to obviating incompatibility problems is to make them from deliberately biodegradable materials. A useful way of classifying the surfaces of long- and medium-term implants is to distinguish those that should ideally completely resist the ad-

Introduction

5

sorption of biomacromolecules (mostly proteins, which constitute the bulk of biomass) from those that are destined to become assimilated as completely as possible with the living tissue, for which the first step is the adsorption of proteins. The first class covers all implants placed in the bloodstream, or in the mouth, or in the presence of some other biofluid (e.g. the contents of the stomach). The second class has the opposite characteristic: in order to promote assimilation with the living tissue it must absorb one or more layers of proteins. This also often applies if the surface has a lubricating role in its contact with the living matter—in this case too, protein or glycoprotein adsorption is a prerequisite. At the same time such implants must also resist the adsorption of proteins that promote the adhesion of and colonization by bacteria. The difficulty of designing a surface able to steer protein adsorption is that living cells in contact with an artificial surface will synthesize and secrete their own proteins; the core task of the designer is somehow to induce the cells to secrete the “right” proteins. Note that the implants of the second class must typically be introduced invasively into the body, and therefore come into contact with the blood before the tissue that is to be their ultimate neighbour, but as far as the blood is concerned, they are only short-term implants. Tissues are usually classified as hard (i.e. mineralized, e.g. bone) or soft. The mechanical properties need to be matched by the mechanical compliance of the implant, but this is not usually a surface property and hence away from the focus of this book. Short-term ‘implants’ are here defined as objects only very briefly in contact with living tissue. This category would therefore include needles, scalpels and other surgical tools. Such objects are not of course really implanted. Many bioanalytical devices should also be included in the short-term category, since the sample (typically a biofluid) may only be in contact with a device for a few seconds. Nevertheless, in view of the required high sensitivity and reliability of such devices, surface characteristics are particularly important for analytical devices. Applications for such systems include blood analysis, cell sorting and characterization, glucose monitoring systems, genetic analysis and general lab-on-a-chip functions. Regardless of the duration of contact, an understanding of protein adsorption is a prerequisite for the successful design of biomedical surfaces. We may anticipate that a profound understanding of the underlying principles will enable both haemocompatible and tissue-compatible surfaces, i.e. both adsorption-resistant and adsorption-promoting, to be designed and fabricated. This knowledge should also permit the fabrication of surfaces having the attribute of resisting bacterial infection. It should always be borne in mind that protein adsorption events may largely or wholly obliterate what-

6

Introduction

ever special, in either a physical or a chemical sense, arrangement of atoms exist on the the surface of a fabricated implant when it is freshly inserted into the body. Only for surfaces destined for extremely short contact, of the order of seconds, with living tissue is the issue of adsorption and important, because there is insufficient time for significant adsorption to take place. The main requirement for such surfaces is simply an extremely low coefficient of friction. Implants have been traditionally considered to be passive inclusions in the living body, and bio-inertness has been striven for. The growing realization that it is essentially impossible for an implant to be truly inert has led to interest in bioactive coatings. The resulting biomedical device is no longer, at best, simply tailored passively (e.g. by finishing it to a certain degree of roughness) to promote some desired response by the host, but is designed to slowly release a hormone or other bioactive substance into its immediate environment over an extended interval (obviously the storage capacity of an implant is finite). The drug delivery particle is a simple example of such a bioactive device, the bioactivity coming from the drug it harbours. Typically the particle also responds to its environment, e.g. one designed to release the drug into the stomach might be coated with impermeable calcium carbonate, which is only dissolved away in the very acidic milieu of the stomach, whereupon it releases its contents. Another example of a bioactive material is a coating prepared with specific biorecognition elements (e.g. short oligopeptide sequences derived from cell surface adhesion proteins) that steer the primary bioresponse to the implant immediately following its insertion into the body. One may conceive a more advanced form of bioactivity, not yet realized in practice, that would aim to steer the metabolism of the bio-environment in which it is in contact in order to make it produce what is necessary to optimize biocompatibility. If nanotechnology succeeds in packing sufficient function into a minute volume, even a miniature factory to make the required drug from available raw materials becomes conceivable. As an objective measure of the relative importance of different body parts and devices, Tables 1 and 2 document the numbers of their occurrences in the very comprehensive (almost 800 A4 pages) list of prostheses issued by the Australian Health Insurance Association (AHIA).

The nature of the bio-environment

Introduction

7

Rank

Body part

number of occurrences

1

hip, pelvic, sacral

764

2

femur, femoral

751

3

knee, patella

704

4

tibia, tibial

575

5

spine, spinal

367

6

(endo)vascular

356

7

heart, cardiac etc.a

271

8

mandible, mandibular

215

9

cervical

192

10

humerus, humeral

142

11

hand, wrist

126

12

shoulder

124

13

ear, ossicular

118

14

(craniomaxillo)facial

96

15

thoracic

92

16

nasal, nose

54

Generic terms 1

bone, orthopaedic

1265

Table 1: Table of the relative importance (ranked by numbers of occurrences of the listed words) of body parts improved by prostheses. Source: Prostheses List, Australian Health Insurance Association (AHIA), 2005. a Also: aortic, atrial, periarticular, ventricular.

8

Introduction

Rank

Device

number of occurrences

1

plate(s)

2434

2

screw(s)

1715

3

stent(s)

533

4

stem(s)

524

5

catheter(s)

351

6

acetabular

252

7

valve

173

8

suture

167

9

pacemaker

79

Generic terms 1

(endo)prosthesis, prostheses

2

fixation

1302 850

Table 2: Table of the relative importance (ranked by numbers of occurrences of the listed words) of implanted devices. Source: Prostheses List, Australian Health Insurance Association (AHIA), 2005.

Introduction

9

The environment of a biomedical implant is very different from that of other, more familiar engineering ones. With their typically warm (37 ◦ C), aqueous and salty nature, the milieux encountered by biomedical materials can be aggressive and give rise to both corrosion- and wear-related problems. From this viewpoint, the most important characteristics are chloride content, dissolved oxygen level and pH. Apart from that, the bio-environment is a very complex medium containing dozens of small molecules, some of which may be ions, and hundreds of different kinds of biopolymers, mainly proteins. Water is invariably the dominant chemical species, but the actual nature of the water, in particular its hydrogen bonding pattern, will depend significantly on the other species present. Interfacial water in particular is likely to be very different from the bulk.3 Cells are present either as a confluent mass, as in the case of tissue, which may be soft (e.g. skin, tendon, pericardium, cornea) or hard (e.g. bone, dentine, cuticle), or as isolated bodies circulating in the aqueous biofluids, such as the blood. As living objects, the cells are more-or-less active, modulating their environment according to the stimuli they receive. One of the most important of these modifications is the coating of an artificial implant with a layer of proteins synthesized and exuded by the cell.

The selection of materials and their attributes The great barrier to the rational design of biomedical surfaces is that while the desired functional attributes are clear enough, e.g. absence of protein adsorption above all for an implant in the bloodstream, or promotion of cell adhesion for an implant in contact with tissue, the physical and chemical attributes with which the surface should be endowed in order for it to achieve those functional attributes still remain largely mysterious, and the work of the designer must therefore inevitably be based on past experience, i.e. empirical knowledge. The repertoire of materials used in biomedical applications (Table 3) is currently quite small, not least because of the necessity of exhaustively testing any new material. One notices the overwhelming popularity of titanium; and at the other end of the spectrum, the very limited use made so far of composites. 3

P. Mentr´e, Interfacial water: a modulator of biological activity, J. Biol. Phys. Chem., 4 (2004) 115–123; P.M. Wiggins, Enzyme reactions and two-state water, J. Biol. Phys. Chem., 2 (2002) 25–37.

10

Introduction

Polyfunctionality. Almost every implant destined for the bloodstream will rest against tissue, and conversely nearly every tissue is vascularized, and hence tissue implants will be in contact with blood (and even if not for long periods, mostly they must at any rate be introduced through a surgical procedure, and hence will be in contact, albeit briefly, with the blood during insertion into the body). Titanium is durable and biocompatible, and is commonly used in heart valves, pacemakers, artificial hips and joints, dental implants and surgical equipment. This metal is frequently alloyed; a typical alloy is Ti-6Al-4V, usually given a TiN coating to overcome the problem of the cytotoxicity of vanadium. NiTiNOL is a superelastic nickel-titanium alloy that is biocompatible, corrosion resistant, and cytocompatible; its shape-memory and elasticity make it a popular choice among manufacturers of implantable devices; vascular stents made from it can be bent to facilitate their insertion into the body with an endoscope and then returned to the prescribed shape by heating to a certain temperature; tantalum is a less usual choice but also offers biocompatibility, durability, and corrosion resistance, and platinum is used where extreme durability is needed. Among the other metals, stainless steel offers good corrosion resistance, formability, and reasonable fatigue resistance. 316L stainless steel is popular; it contains Cr, Mo, Ni, Mn, C, S, Si, P and Fe. Cobalt-chromium alloys (Co-Cr-Mo, Co-Cr-Mo-Ni, Co-Cr-MoNi-Fe, Co-Cr-Mo-Ni-W-Fe and Co-Cr-W-Ni) offer good corrosion and wear resistance. Ceramics include alumina (Al2 O3 , typically containing a trace (< 0.3 wt% of magnesia (MgO)), zirconia (ZrO2 ), hydroxyapatite (Ca10 (PO4 )6 (OH)2 ), tricalcium phosphate (Ca3 (PO4 )2 ), so-called “bioglass” (Na2 O(CaO)(P2 O3 )(SiO2 )), and calcium aluminate (Ca(Al2 O4 )). Presently-used biomedical composites almost invariably consist of polymers (PEEK, UHMWPE, PMMA) incorporating carbon fibres. Table 4 gives a selection of the typical applications of the various biomedical materials. The potential interactions of the surface of material with its environment are ultimately determined by energetics. The interfacial energetics of a surface can be compactly expressed in terms of surface tensions,4 which are in turn determined by the atomic and molecular composition, which also of course determine the chemical behaviour. This will be explained in more detail in Chapter 2, but it is already useful to note that the interfacial free 4

Fairly extensive tabulations of the surface tensions of a wide range of solids, including metal oxides, organic polymers, and proteins are already available (e.g. Table 2.1.1 in Chapter 2). They provide an important resource for initial design considerations regarding biomedical surfaces.

Introduction

11

Rank

Material

number of occurrences

1

titanium

2492

2

steel

607

3

chromium

391

4

polyethylene (PE), UHMWPE

389

5

silicone

375

6

cobalt

235

7

PTFE, Teflon, fluoroplastic

220

8

molybdenum

216

9

vitallium

198

10

polypropylene (PP)

160

11

PMMA (polymethylmethacrylate), acrylic

111

12

polyester

107

13

NiTiNOL

101

14

polyurethane (PU)

72

15

platinum

58

16

premilenea

55

17

aluminium, alumina

52

18

carbon

51

19

rubber

49

20

hydroxyapatite

45

21

silicon

42

22

PLLA

39

23

PEEK (polyethyletherketone)

32

23

tantalum

18

Generic terms 1

plastic(s), (co)polymer

304

2

alloy

302

3

metal

190

4

ceramic

157

5

elastomer

40

6

composite

32

Table 3: Table of the relative importance (ranked by numbers of occurrences of the listed words) of the materials used in implanted devices. Source: Prostheses List, Australian Health Insurance Association (AHIA), 2005. a A lightweight polypropylene mesh.

12

Introduction Application

Material

Replacement joints

titanium, stainless steel, PE

Plate (bone fracture repair)

stainless steel, Co-Cr

Cement (for bones)

PMMA

Filler (for bones)

hydroxyapatite

Breast enlargement

silicone

Replacement tendons and ligaments

PTFE

Stents

NiTiNOL, stainless steel

Replacement heart

polyurethane

Replacement cochlea

platinum

Table 4: Table of typical applications of biomedical materials. energy of an object (e.g. a macromolecular particle) made from material 3 with a substrate made from material 1 in the presence of a liquid medium 2 is given by ∆G123 = ∆G22 + ∆G13 − ∆G12 − ∆G23 , (1) i.e. the interfacial free energy is the sum of the cohesive energy of the liquid and the direct energy of interaction between the surface and the particle in the absence of liquid, minus the solvation energies of surface (adsorbent) and particle (adsorbate). Each term in equation (1) is itself the linear sum of the Lifshitz-van der Waals, electrostatic, and Lewis acid/base interactions. In aqueous systems, the main manifestation of the Lewis acid/base interaction, (also known as electron donor-acceptor interaction) is hydrogen bonding. This comprises the dominant contribution—quite possibly as much as 80 or 90%—to the total interaction energy between say a protein and a surface. Van der Waals interactions are ubiquitous but weak, and although many biomedical surfaces are electrostatically charged, typically because of protonation or deprotonation of surface groups, these charges are strongly shielded under usual physiological conditions and the electrostatic interactions are therefore weak. The hydrophobicity of a surface has often been considered to be a reliable indicator of its biocompatibility, but despite his frequent discussion in the literature, the matter gives rise to much confusion. Materials such as PTFE that have very low surface energies (they are in fact both hydrophobic and oleophobic) indeed tend to make the second and third terms on the right-

Introduction

13

hand side of equation (1) small, but since these two terms are of opposite sign the net effect will tend to cancel out and be close to zero. The choice of material will have no effect at all on the first and the last terms. The “hydrophobic force” is simply a consequence of a very large and negative cohesive energy of water (i.e. the first term on the right hand side with 2 = water). In the absence of all other interactions, any two materials 1 and 3 will tend to stick together (i.e. have a large and negative ∆G123 ) simply by virtue of this fact. Hydrophobic attraction can only be opposed if “hydrophilic repulsion” (the last two terms on the right-hand side) is sufficiently strong to overcome the cohesive energy of water. A surface like PTFE attracts water weakly like most other surfaces by the van der Waals interactions. Proteins cannot form hydrogen bonds with such surfaces because they are devoid of electron donors and electron acceptors. The last term of equation (1) can however be made even smaller if a protein undergoes eversion, i.e. it is conformationally rearranged to expose its hydrophobic core to the PTFE surface. This probably happens in the case of PTFE exposed to blood: albumin, the most abundant blood protein, weakly adsorbs and then probably everts, assuming a non-native confirmation. Nevertheless, adsorbed albumin is believed to be a rather bioincompatible surface coating. This is somewhat surprising because conformationally altered proteins tend to be immunogenic, since they have become non-native, i.e. non-self, and a recognized as such by the immune system. The absence of significant adsorption of other proteins to adsorbed albumin, or the absence of cell attachment, as ascertained in laboratory tests may be very far from representative of the real situation once the albumin-coated surface is inserted into the human body, because that environment is far more complex, possibly containing hundreds of different proteins as well as a many small molecules. The longer the implant resides in the body, the greater the discrepancy between that environment and that of the laboratory. More careful and detailed considerations, as will be introduced in the subsequent chapters, are needed to properly evaluate biocompatibility. Much of the earlier work was based on very speculative notions, which have often persisted even in the face of clear experimental refutations. The situation now is however much more favourable because of the explosion of high resolution techniques such as “molecular microscopy”, which enable biomedical surfaces to be probed on the nanoscale. Hydrophobicity can be formally defined by generating a new equation from (1), by replacing material 3 by another sample of material 1, to get ∆G121 = ∆G22 + ∆G11 − 2∆G12 .

(2)

If ∆G121 > 0, material 1 (with 2 = water) is hydrophilic, i.e. two samples

14

Introduction

of the material placed in water will not aggregate; if ∆G121 < 0, material 1 is hydrophobic (because the interfacial hydration represented by ∆G12 is insufficient to counter the strong attractive force originating from the cohesive strength of water ∆G22 . Clearly this approach can be generalized to provide definitions of solvophobicity and solvophilicity, but water is obviously the most important solvent for any biomedical application. In an actual application, the biomedical surface engineer has control over the selection and processing of material 1, but has no control over material 2 (the biofluid), nor over material 3 (the living matter). Nevertheless, this does permit control over two of the terms in equation (1), namely ∆G12 and ∆G13 , and these should be given in material selection charts for biomedical applications. Nevertheless, both ∆G23 and the very large and negative ∆G22 of water remain uncontrollable. Even this relatively simple picture is likely to be inaccurate, however, because of the tendency of any material inserted into a living body to become modified by the secretions etc. of the body. The main physical attribute of a surface is its roughness. The traditional approach adopted by manufacturers has been to finish the surfaces of devices destined to be placed into the bloodstream as smooth as possible (in order to minimize protein adsorption, platelet attachment etc.), and to deliberately roughen the surfaces of devices destined for assimilation into the body in order to provide a grip for the cells. Systematic data on the influence of roughness—at all relevant length scales—on the responses evoked in biological milieux is still largely lacking, however. For proteins considered as material 3, the above considerations, implying averaged surface energies, are only a first, or even zeroth, approximation, because of the significant chemical heterogeneity of protein surfaces. Proteins are made up of folded amino acid chains, and the amino acid side chains (residues) may be polar or apolar, and among the polar residues there are those of hydrogen bond-donating (i.e. electron accepting) and hydrogen bond-accepting (i.e. electron donating) character, and some residues can be electrostatically positively or negatively charged through the accumulation of depletion of hydrogen ions. The old idea was simply that complementarities between the chemical functionalities of the protein (the adsorbate) and the biomedical surface (the adsorbent) determined adsorption (adhesion). Current debates about its validity are somewhat reminiscent of the old discussions in medicine regarding contraria contrariis curantur versus similia similibus curantur. Thus, positively ionized groups (e.g. amines) on the surface should strongly adsorb proteins with an excess of, say, aspartic acid residues, and patches of apolar residues should strongly adhere to apolar (“hydrophobic”) surfaces (such as polyethylene). Recognition that the ionic strength of most biological milieux

Introduction

15

strongly shields electrostatic charges resulted in an improved version of this approach, based on electron donors and acceptors rather than proton donors and acceptors, i.e. on Lewis acids and bases rather than on Brønsted acids and bases. But this failed to take into account that these interactions take place not only in the presence of dissolved ions but also in the presence of an overwhelming excess of water, which is both an excellent Lewis acid and an excellent Lewis base. A key concept is therefore the degree to which hydrogen bonds can be “wrapped” or desolvated by the presence of apolar chemical groups in their vicinity.5 And all these considerations only concern the initial approach of the protein to the surface: as soon as the protein is able to reside for some time actually on the surface, the chemical groups of the surface are able to compete with the groups on the protein to form new bonds, at the expense of the intramolecular conformation-determining bonds of protein, possibly resulting in denaturation. the vicinity of the backbone hydrogen bonds. There is a therefore a complex interplay of intramolecular and intermolecular (i.e. protein-surface) bonding that needs to be considered when attempting to predict the fate of a protein arriving in the vicinity of a biomedical surface. This will be discussed in more detail in the subsequent chapters. Since cells have many proteins on their surfaces, the interaction of cells with artificial surfaces is also largely determined by protein-surface interactions (independently of the fact that cells on a substrate typically synthesize and release proteins, the so-called “microexudates”, that are then adsorbed on the substrate). The cell surface proteins are typically so-called “membrane proteins”, i.e. embedded in the lipid bilayer membrane enveloping the cell. The parts of the polypeptide chain that are within the membrane obviously do not come into contact with the biomedical surfaces, but most of these proteins have large external domains similar to those of soluble globular proteins, except for the anchoring points where the polypeptide of the chain enters the membrane. Biomedical materials have traditionally been selected based on their bulk characteristics—properties necessary to ensure that they can be shaped inconvenient manufacturing processes, and retain mechanical integrity once in the body. However, the surface properties may not always be particularly biocompatible, and therefore a useful and important design strategy is to make use of surface engineering, which aims to enhance corrosion and wear resistance, antibacterial characteristics and tissue compatibility (including suppressing any adverse immune response). In summary, therefore, biocompatibility is an active, dynamical process 5

A dehydron is an underwrapped or under-desolvated hydrogen bond.

16

Introduction

involving a two-way exchange of information between the implant and the host that needs to be managed in order to be achieved. There are of course constraints imposed by the required bulk attributes, and practicability of manufacture etc., but if necessary a bulk material can be surface engineered or coated. Overarching definitions of biocompatibility need to be translated into specific physico-chemical attributes of the material. Even if a more precise description of biocompatibility in a given situation can be made—e.g. a biocompatible stent needs to resist protein adsorption, for example—the relationship between protein adsorption resistance (which we can call a functional attribute) and physico-chemical surface attributes such as surface tension and roughness is not yet known in general. Figure 1 sketches out an idealized chain of inference. design Figure 1: The relationship between applications and attributes.

Past experience is clearly shown that attempts to subsume complex surface attributes into a single parameter such as hydrophobicity are inadequate for predicting and managing biocompatibility. Thanks to nanotechnology, materials science stands at the threshold of a renaissance. The possibility of atomic precision in the assembly of materials of possibly arbitrary complexity opens vast new possibilities for the design and fabrication of materials with tailored properties. Nevertheless, these possibilities are not immediately accessible to the designer of a biomedical surface, because while the required functional attributes can sometimes be intelligibly and unambiguously expressed, correlations between those functional attributes and the physical and chemical attributes (cf. Chapter 2) are still largely unknown (cf. Chapter 8). The goal of this book is to pave the way for that renaissance.

Chapter 1 The adsorption and desorption of biomacromolecules An essential feature of the behaviour of artificial materials inserted into most places within a living body is that already after a very short interval they will become covered by proteins and other biopolymers, obliterating any careful two and three-dimensional structuring, both chemical and morphological, that may have been applied to the surface during fabrication. As discussed in the Introduction, the two extremes of the range of biomedical milieux encountered by biomedical devices are the bloodstream, and more-or-less solid tissue. Blood contains a great variety of different proteins. Apart from sharing the attributes of solubility, and an overall excess of dativit`a (see Table 2.1.1), ensuring that they do not clump together (sickle cell haemoglobin is of course a pathological case that no longer fulfils the complete set of solubility criteria, as will be explained in §2.1.3), structurally speaking these different proteins have little in common. In particular, one should note the presence of both compact, globular proteins and fibrous proteins; typically under physiological conditions, the fibrous proteins will also tend to have a compact shape.1 The marginal stability of even the globular proteins with respect to conformational change indicates that slight environmental changes are sufficient to move the system into another basin of attraction, corresponding to aggregation and precipitation. The introduction of an artificial surface into the bloodstream is typically more than a slight environmental change, and in very many cases leads to those buildups of adsorbed insoluble deposits that are so troublesome, both for the correct functioning of the implanted device, 1

Rocco, M., Infusini, E., Daga, M.G., Gogioso, L. and Cuniberti, C. (1987). Models of fibronectin. EMBO J. 6, 2343–2349.

17

18CHAPTER 1. THE ADSORPTION AND DESORPTION OF BIOMACROMOLECULES and for the continued health of the host organism, who is now at risk of embolism (a blocked blood vessel). At the other extreme, artificial surfaces introduced into solid tissue, it is well-established that the cells in intimate contact with an artificial surface synthesize and excrete their own proteins with which they cover the surface, as a precursor to their own subsequent semi-permanent attachment. The conclusion is that protein adsorption cannot be evaded. The ultimate goal of the biomedical surface engineer is therefore to design the surfaces in such a way that control over these adsorption processes is retained. As will become apparent, however, this is a very difficult challenge, which is far from being solved at present.

1.1

Elementary physico-chemical considerations

Regardless of the details of the actual forces involved, one can already make a general statement about the energetics of protein-surface interaction (by ‘protein’ we mean essentially any kind of biological macromolecule, and by ‘surface’ we mean the substrate, assumed to be perfectly planar for the present, on which the macromolecule is adsorbing), as is sketched out in Figure 1.1. The mass action law (MAL) allows us to derive the thermodynamic affinity, characterized by the equilibrium constant K, from the energetics: ∆Gd = RT ln K

(1.1)

where R and T have their usual meanings as the gas constant and the absolute temperature respectively. The same quantity in principle—at least in simple cases) can be determined by the kinetic mass action law (KMAL): Kkin = ka /kd

(1.2)

where the k are the rate coefficients for adsorption and desorption. Supposing that they can both be approximated using Arrhenius-type expressions, we can write: ka = fa exp −∆Ga /(RT ) (1.3) and kd = fd exp −(∆Gd + ∆Ga )/(RT ) .

(1.4)

Substituting these last two equations into (1.2) we then get Kkin = (fa /fd ) exp ∆Gd ,

(1.5)

1.1. ELEMENTARY PHYSICO-CHEMICAL CONSIDERATIONS

19

intEnProf Figure 1.1: Sketch of the interfacial potential energy acting between a compact macromolecule approaching an adsorbent surface. The abscissa gives the separation between the macromolecule and the surface, and the ordinate the potential energy. Far from the surface (and for the physiological situations relevant to this book, that may mean as little as 10–20 nm) there is no perceptible interaction, and the molecule moves with the motion of an unbiased random walk. If it happens to approach the surface, it will typically sense an increasingly repulsive force. As the molecule approaches yet closer, the repulsive force is increasingly counteracted by a shorter range attractive force (this is the quasi-universal Lifshitz-van der Waals interaction), which at very close range (a few nanometres) increasingly dominates, and those proteins that manage to diffuse over the thermal barrier (height ∆Ga ) peaking at the crossover between the long-range repulsive and short range attractive forces fall into a more or less deep potential well, whose depth ∆Gd is determined by the onset of the extremely short-range Born repulsion.

i.e. the same as eqn (1.1), provided fa = fd (for which there is, strictly speaking, no evidence).2 The above considerations are indeed elementary and not really applicable to the adsorption and desorption of macromolecules, especially biomacromolecules, because typically the species are changing (evolving) during the adsorption and desorption process. During the approach of the molecule to the surface, the environment is changing both in terms of water structure and in ionic composition. This can lead to surprising—at first sight—results; for example, an excess of protons attracted to a strongly negatively charged surface (i.e. it has a low ionic strength) may reverse the net electrostatic charge of a protein. Once the molecule arrives at the surface (i.e. finds itself in the deep potential well in Figure 1.1, it may well undergo more or less drastic conformational 2

Another difficulty that emerges from these considerations is that the usual units of the adsorption and desorption rate coefficients are different ([L][T]−1 and [T]−1 respectively), meaning that Kkin in eqn (1.2) is not dimensionless—it has the dimensions of length. One way round this difficulty is to completely separate the transport towards the surface from the action of surmounting the energy barrier (see Figure 1.1), and suppose that the coefficient ka applies to a pure first order process.

20CHAPTER 1. THE ADSORPTION AND DESORPTION OF BIOMACROMOLECULES changes. The primary driver for these conformational changes comes from the exchange of the weak intramolecular bonds (hydrogen bonds) that hold the protein structure together for weak bonds of similar type between the molecule and the substrate on which it is adsorbing. It is very possible that there is no significant change of enthalpy H upon adsorption. As soon as that exchange has taken place, many more conformational possibilities are open to the molecule (Figure 1.2), and its entropy S increases, and, according to the relation ∆G = ∆H − T ∆S , (1.6) the overall free energy for the process will be negative. surfDenat Figure 1.2: Sketch of the adsorption process for a typical globular protein (after Fern´ andez and Ramsden, 2002). Phenomenologically, it is very well known that typically proteins undergo marked hysteresis upon adsorption/desorption. We consider here only adsorption from solution—gas phase adsorption is scarcely of relevance. When a surface is exposed to a solution of proteins they rapidly adsorb, but when the proteins are then removed from the solution (or the solution is replaced by pure solvent) desorption is extremely slow. Intensive efforts have been undertaken to determine more precisely the exact nature of the conformational changes (the experiments are difficult because the measurements should be carried out in situ). In some particular cases, that may perhaps be considered to be typical, although the subset of proteins that have been investigated (serum albumin and lysozyme seem to be the most popular choices) can in no sense be considered to be representative of the tens of thousands of different proteins interacting with surfaces in human body), there is evidence for adsorption-induced loss of protein secondary structure, which is relatively easy to detect, more-or-less drastic changes in enzyme activity following adsorption, and history-dependent desorption. Due to this structural evolution, the problem formally becomes one belonging to complexity science. At the time of writing, adequate (in the sense of fully comprehensive) models for the process have still not been developed, but work is in progress. The consequences of this conformational evolution are dramatic. Clearly the system (surface + liquid environment + protein) is non-ergodic, with the consequence that the ensemble average at a particular epoch is not equivalent to the time average of a particular molecule. This is indeed already familiar in glassy systems, and some of the previous work on glassy relaxation

1.2. MONOLAYER ADSORPTION

21

is applicable to protein adsorption-desorption. This non-ergodicity should constantly be borne in mind. The models that will be described in what follows are typically applicable to a restricted subset of phenomena. Figure 1.3 attempts to capture the overall process graphically. adsEv Figure 1.3: Sketch of the events associated with the approach, arrival and residence of a protein to, at and on a biomedical surface.

1.2

Monolayer adsorption

In this section, we shall model the adsorbing protein as an immutable hard sphere adhering irreversibly to the surface. This relatively simple model already leads to nontrivial adsorption behaviour, and as has been already demonstrated experimentally, is well-applicable to certain globular proteins such as the cytochromes, which are rather strongly internally bonded and do not tend to undergo denaturation upon adsorption. The main important difference between proteins adsorbing on a biomedical surfaces and the familiar Langmuir-type adsorption of gases on the metal is that whereas in the latter case the adsorbate is adsorbed on discrete, welldefined adsorbent sites smaller than the adsorbate atom or molecule, the protein is much larger than any individual point of anchorage. Typically a protein has many sites on its molecular surface able to weakly bond (e.g. via hydrogen bonds—details of the bonding will be left until Chapter 2) with the substrate surface, and the substrate is also densely covered with bonding partner sites. Hence the protein is essentially adsorbing on a continuum.3 The incoming protein will occupy a certain area a when it lands on the surface. Note that the actual position is selected at random. Even under conditions of laminar or turbulent flow, in the vicinity of the surface friction strongly retards convection, and macromolecule transport is by diffusion.4 The fraction φ of the surface available for adsorption, initially of course equal to unity, is thereby diminished. In the Langmuir model of adsorption, if the 3

The following argument also holds for proteins (e.g. antibodies) adsorbing on to a carpet of receptors (e.g. antigens), provided the receptors smaller than the incoming ligand. This is the situation encountered in many biochemical assays, including biosensors. However, it will not be considered in more detail here since this is not strictly speaking part of the topic of this book. 4 The combination is called ‘convective diffusion’ by Levich.

22CHAPTER 1. THE ADSORPTION AND DESORPTION OF BIOMACROMOLECULES receptor sites had an area exactly equal to or greater than a, we would have the relation φ = 1 − θ = 1 − νa . (1.7) This expression is not however correct if the surface is a continuum, lacking discrete sites. The deposition of a sphere creates an exclusion zone around it, the meaning of which is that the centre of a subsequently arriving particle cannot land within the exclusion zone. Clearly the exclusion zone has twice the diameter of the particle, and hence we have φ = 1 − 4θ = 1 − 4νa .

(1.8)

It will be obvious that as the surface becomes filled up, we thereby overcompensate for the loss of available area for adsorption and must add back terms proportional to θ2 for the overlap of two exclusion zones, and proportional to θ3 for the overlap of three exclusion zones: φ = 1 − 4θ + b2 θ2 + b3 θ3 + . . . .

(1.9)

The series continues until the sixth power, and then stops, since a maximum of six spheres can be arranged around another one on the plane, but for practical purposes it is sufficient to continue until the third power. This is the model of random sequential addition, usually abbreviated as RSA, and the coefficients b2 and b3 have been worked out exactly for spheres, as well as for some other simple shapes. The model strongly resembles the virial expansion for an equilibrated arrangement of spheres on a plane, first explored by Widom. Of considerable practical importance is the fact that in two dimensions, ψ becomes zero at a jamming limit θJ of a little over 50% for spheres.5 The parameters θ and ψ are crucial to an understanding of adsorption phenomena. The fractional surface coverage θ directly monitors the amount of protein on the surface; in a practical experiments, this might be measured as mass M or number of moles M per unit area, in which case we have θ = M a/m and θ = M aNL respectively, where a is the area occupied on the surface by one molecule, m is the mass of one molecule, and NL is the Loschmidt (Avogadro’s) number. On the other hand the rate of deposition dθ/dt is directly proportional to φ, the constant of proportionality being as a rule the product of a rate coefficient and the bulk concentration. If the adsorption mechanism is established as being RSA, then a single parameter suffices to characterize it, namely a. We can therefore write (dθ/dt)RSA = ka cφRSA 5

The value of θJ is slightly different for more elongated shapes.

(1.10)

1.2. MONOLAYER ADSORPTION

23

as the general kinetic equation, where ka is the adsorption rate coefficient, and c equals, or approximately equals, the bulk concentration cb . A quite different mode of adsorption is called ballistic deposition (BD). In this case the incoming objects have a vanishingly small probability of sticking to the surface (in practice some imperfection in the surface will become an initial adhesion site). Particles landing on a vacant area of an already partly occupied surface will migrate to the edge of an existing deposit and there adhere, and particles landing on an actual deposit will similarly migrate to the edge of the deposit and there adhere. A more realistic model that avoids the difficulty associated with initiating BD combines strict RSA with the BD mechanism, introducing the parameter j φGBD = 1 + f1 (j)θ + f2 (j)θ2 + f3 (j)θ3

(1.11)

f1 = 4(j − 1) ,

(1.12)

f2 = 3.808 − 0.180j − 3.128j 2

(1.13)

f3 = 1.407 + 4.679j − 25.58j 2 + 8.550j 3 .

(1.14)

where

and The full derivation of the coefficients, appropriate for spherical particles, is given by Tarjus et al. (1994).6 Simple as the RSA model is, it provides an astonishingly good description of many observed protein deposition processes. One reason for the good agreement between theory and experiment is the fortuitous cancelling out of the correlating effects of successive adsorption attempts, and the anticorrelating effects of the hydrodynamic friction acting on the particle motion in the direction perpendicular to the surface.7 When the first, and possibly some subsequent adsorption attempts, are rejected because the area chosen is already occupied, in reality the particle does not simply disappear back into the bulk and try again somewhere else, which is what is usually done in simulations of RSA: the Rabinovich (cage) effect ensures that successive attempts are typically in the immediate vicinity of the initial one, and this has also been observed in explicit simulations;8 This is the origin of the strong correlations between the positions of successive adsorption attempts. 6

Tarjus, G., Viot, P. Choi, H.S. and Talbot, J. Phys. Rev. E 49 (1994) and 3239. Bafaluy, J., Senger, B., Voegel, J.-C. and Schaaf, P. (1993). Effect of hydrodynamic interactions on the distribution of adhering Brownian particles. Phys. Rev. Lett. 70, 623–626. 8 Luthi, P.O., Ramsden, J.J. and Chopard, B. (1997). The role of diffusion in irreversible deposition. Phys. Rev. E 55, 3111–3115. 7

24CHAPTER 1. THE ADSORPTION AND DESORPTION OF BIOMACROMOLECULES The hydrodynamic friction has a randomizing effect, that just happens to exactly compensate for the correlating effect. On the other hand, cases are also known where the lateral surface mobility of the adsorbing protein, and its propensity to form clusters, is so high that in the adsorption kinetics there is no trace of the initial random sequential deposition, and the kinetic equation collapses to a simple Langmuir expression φ ∼ 1 − θ. The mobility of proteins on solid surfaces seems to be a common phenomenon; and was already experimentally demonstrated (for albumin on glass) in 1980.9 The surface denaturation described above (see Figure 1.2) implies that the adsorbed protein expands the area it occupies on the surface. Although evidence for such expansion can be obtained by studying the variation of the area a per adsorbed molecule with bulk concentration, on the premiss that at higher bulk concentrations the adsorption of many particles takes place too rapidly to allow time for the adsorbed particles to expand, a more exact treatment takes explicit account of the requirement for space into which the expansion could take place: analogously to the function φ, we create a function Ψ that gives the fraction of the surface available for expansion, and the kinetics become dθ/dt = ka cφ − kdenat θΨ

(1.15)

dθdenat /dt = kdenat θΨ

(1.16)

and

where θdenat is a fraction of the surface occupied by denatured proteins, and kdenat is the rate coefficient for surface denaturation. If surface denaturation with expansion is taking place, then the jamming limit is likely to be bulk concentration-dependant, depending on the relative magnitudes of the two rate coefficients ka and kdenat . If the expansion is rather fast, the surface will rapidly become jammed with respect to the denatured form, but can then still accept further native molecules, which cannot expand; if on the other hand the expansion is rather slow (or the bulk concentration is very high), the surface will become jammed with respect to the native form, and very little denaturation can take place. This suggests a useful strategy if it is desired to prepare an active surface coating with a protein that becomes inactivated by being denatured. It should not be supposed that all proteins expand when denatured. The does indeed appear to be the most common type of behaviour, that there 9

Michaeli, I., D., Absolom, R. and van Oss, C. J. (1980). Diffusion of adsorbed protein within the plane of adsorption. J. Colloid Interface Sci. 77, 586–587.

1.2. MONOLAYER ADSORPTION

25

are “natively expanded” proteins,10 which cannot further expand; any conformational change must therefore lead to their contraction.

1.2.1

Effect of flow rate

In the preceding development, the actual number of adsorption attempts per unit area per unit time have been subsumed into the concentration c without inquiring too closely about details. In the practical case of adsorption of blood proteins onto a cylindrical stent in a blood vessel, we have Poiseulle flow, i.e. the characteristic parabolic profile of liquid velocities, zero at the walls of the tube (cylinder, i.e. the blood vessel), and at a maximum in the centre. Therefore, at some point within the tube diffusion of dissolved species will be faster than convection. This is illustrated in Figure 1.4. Beyond the diffusion boundary layer, i.e. in the central portion of the tube, the concentration of dissolved adsorbent is constant and equal to the bulk concentration cb . At the surface of the adsorbate the concentration is zero. The energy barrier (Figure 1.1) is at a distance of the order of 1 nm from the surface, and lies between the dotted line on Figure 1.4 and the surface proper. If there were no energy barrier, they would be a simple diffusion gradient cb /δh driving the adsorbing species onto the surface, and the rate of deposition would be Dcb /δh , where D is the diffusion coefficient of the adsorbing species. Due to the energy barrier, there is accumulation of material in front of it. Let us suppose that the concentration of material immediately before the energy barrier is only slowly varying (compared with other relevant processes), and has a value of c1 . Then we can write dc1 /dt ≈ D(cb − c1 )/δh − ka c1

(1.17)

making use of the fact that δh  the distance of the pre-energy barrier zone from the surface, yielding c1 = cb /(1 + ka δh /D) .

(1.18)

If therefore δh ka  D, then c1 ≈ cb . Expressions for δh are given by Levich. A useful one is δ = (3/2)2/3 (DxRA/F)1/3 , (1.19) where F is the volumetric flow rate fluid through the cylinder (the blood vessel), R is the radius of the blood vessel and A its cross-sectional area, and x is the distance from its beginning. The above treatment assumes that there is no impediment to the adsorption of any species that surmounts the energy

26CHAPTER 1. THE ADSORPTION AND DESORPTION OF BIOMACROMOLECULES conDiff Figure 1.4: Convective diffusion in a blood vessel. v is the velocity of the liquid. The diffusion boundary is indicated by the dashed lines. The dotted line indicates a hypothetical layer whose thickness is of molecular dimensions.

barrier, i.e. φ = 1. If that is not the case, a corresponding modification has to be made. When comparing experimental results from different flow r´egimes, it is particularly important to ensure that one is comparing like with like, and that appropriate normalizations are carried out. The theoretical analysis of multicomponent adsorption is complicated by the fact that each species has its own value of the hydrodynamic diffusion boundary thickness, which depends not only on the flow rate but also on the diffusion coefficient. Many soluble proteins have rather similar values of D however, and besides δh only depends on the cube root of D. This becomes an issue only when comparing objects of very different sizes, such as proteins with bacteria or eucaryotic cells.

1.3

Desorption

The energy profile sketched in Figure 1.1 implies (dθ/dt)des = −kd θ

(1.20)

which has the familiar exponential decay as a solution θ(t) = θ0 exp(−kd t)

(1.21)

where θ0 is the surface coverage at the start of desorption. As already pointed out in §1.1 however, except in a few special cases the corresponding behaviour is not observed experimentally; the desorption kinetics are strongly nonexponential. In view of processes such as those illustrated in Figure 1.2, this is hardly surprising. Indeed, with regard to the multiplicity of possible mechanisms that could lead to non-exponential desorption kinetics, the best approach to describe the phenomenon is in terms of a memory function. Let Q(t, t1 ) be the memory 10

D´er et al., 2007.

1.4. ADSORBED PROTEIN EXCHANGE

27

kernel. It gives the probability that a molecule adsorbed at epoch t1 will still be adsorbed at a later epoch t. Then11 Z

t

φ(t1 )Q(t, t1 )dt1 .

θ(t) = ka c

(1.22)

0

ka and c have been placed outside the integral, but in the most general case, they could also be time-dependent. If there is no desorption (irreversible adsorption), then obviously Q = 1; Q = exp(−kd t) corresponds to the complete absence of memory (a Poisson process). The memory function can be interpreted in terms of more explicit models. One that is rather popular in the protein adsorption-desorption literature, and which has often been invoked to discuss experimental results, conceives the protein initially adsorbing in its native (i.e. dissolved) conformation, in which it can desorb without memory, i.e. according to eqn (1.20), but also capable of undergoing a transition to a conformationally altered state, in which it can desorb much more slowly, or not at all on practical laboratory timescales.12 If Q does not depend on θ, then it may be written as a function of the time difference t − t1 , but this also appears to be a special case, and in general the desorption probability of a given molecule does depend on the surface coverage, i.e. even in the simple equation (1.20) kd should be a function of θ. Once all these aspects are taken into account (including the fact that the conformationally altered state will probably occupy more area on the surface than the native state), it is apparent that quite a rich behaviour is possible even for a monolayer formed from a single component in the bulk.

1.4

Adsorbed protein exchange

Under conditions in which the desorption of a fully adsorbed protein monolayer (i.e. at the jamming limit) is known to be extremely slow, experiments in which such a layer, stable in contact with pure solvent, is then exposed to a solution of an almost identical protein, differing only in that it carries a label, typically radioactive iodine, have shown the acquisition of labelled proteins by the layer, and conversely a fully adsorbed monolayer of labelled protein, stable in contact with pure solvent, releases label (radioactivity) into the solvent when exposed to a solution of unlabelled protein, without any change in the total adsorbed mass, in other words the reaction (denoting 11

Talbot, 1996. Note that the criterion for the establishment of equilibrium is that the memory function goes to zero at long times. 12

28CHAPTER 1. THE ADSORPTION AND DESORPTION OF BIOMACROMOLECULES protein by P and the presence of a label by a star) Pads + P∗soln P∗ads + Psol

(1.23)

takes place. Intriguingly, the kinetics are comparable to, or even faster than, the kinetics of the adsorption of the protein in their layer. No molecular mechanism has been advanced for process, and since the label is, as mentioned, typically radioactive iodine, the possibility cannot be excluded that it is simply the label that is being exchanged, rather than the protein itself.

1.5

Multilayer adsorption

The model of monolayer adsorption has turned to be a very useful one for elucidating the basic mechanisms of protein adsorption and desorption. In completely natural biological systems, it applies to most of the enzyme-catalysed reactions taking place at membrane surfaces, and is especially relevant for understanding the performance of biomedical implants in the bloodstream (see Chapter 3). Structural protein deposits in biological systems are typically composed of multilayers. The basic models developed in the preceding sections can also be applied, in simplified form since at least to a first approximation the available area function ψ is always equal to unity. The component protein (e.g. laminin, fibronectin) typically belongs to the class of fibrous proteins, which have elongated, indeterminate structures even in solution, and hence the conformational transitions illustrated in Figure 1.2 are not of such importance.13 On the other hand, proteins in layers deposited early become overlaid with later deposits, which must be removed before they can be desorbed. Such a sequential hierarchy of desorption gives rise to Kohlrausch or stretched exponential behaviour, i.e. eqn (1.21) must be modified (“stretched”) to become θ(t) = θ0 exp(−kd t)β

(1.24)

where the Kohlrausch exponent 0 < β < 1. This behaviour has recently been experimentally verified.14 . Considerable structural richness is possible in such a multilayer, even if it is composed of a single protein component, e.g. it may show form anisotropy, anomalous surface roughness, etc. 13

Note that the structure of many fibrous proteins can be modelled as a daisy chain or a pearl necklace: a string of compact globules joined by short loops (Rocco et al., 1987). The compact globules probably keep their structural integrity even when the protein is adsorbed. 14 McColl et al., 2007).

1.6. MULTICOMPONENT ADSORPTION

1.6

29

Multicomponent adsorption

Real blood contains hundreds of different proteins. It has that intermediate level of variety that is very difficult to handle. The simple case of a binary mixture of large and small particles can be tackled: each particle size has its own corresponding available area function, and the two are coupled. Roughly speaking, the larger particles simply reduce the available area without much affecting the spatial correlations of the smaller particles. This problem is related to adsorption with the post-conformational transition (cf. equations (1.15) and (1.16)). This approach quickly becomes unworkable as the number of components increases. On the other hand, the variety is not so large that it can be treated statistically. There is a real difficulty in that routine identification of all the adsorbed species cannot currently be undertaken, and therefore the experimental basis for deciding whether some anomaly in the adsorption observed with an implant placed in the blood is due to an unusual adsorbed protein composition (which might simply be the absence of one protein type) is lacking. There are certainly no grounds for supposing that the composition of the adsorbed protein layer mimics the bulk protein composition. Clearly, there are fascinating possibilities for the assembly of complex three-dimensional structures from cell to starting material. High resolution experimental observations on the accumulation of adsorbed adlayer from blood plasma have revealed that the composition (as determined by the binding of selected antibodies) of the layer is constantly evolving, confirming the well established Vroman effect, and have furthermore shown that the evolution depends on physical variables such as flow rate that have not been hitherto considered to qualitatively affect thin film structure.

1.7

Adsorption onto heterogeneous surfaces

Composite materials are currently enjoying a renaissance, driven by nanotechnology. This implies that many future biomedical materials will have chemically heterogeneous surfaces. We have already pointed out that the surfaces of proteins are heterogeneous—at a length scale of the order of 1 nm, the light of an amino acid residue. Atomic heterogeneity, as exemplified for example by an alloy, the constituents of which form a solid solution with each other, might not therefore be significant, and macroscopic heterogeneity, with micrometre-sized islands, might well be treated as the area-weighted mean of the pure constituents, but if the heterogeneity is of a scale of the order of that of the protein, “anomalous” effects might well arise.

30CHAPTER 1. THE ADSORPTION AND DESORPTION OF BIOMACROMOLECULES In the a living cell, the proteins themselves are distributed heterogeneously on the surface, suggesting that (depending on the size of the cell) heterogeneity at a scale of the order of 10–100 nm might result in “anomalous” behaviour in cell adhesion.

1.8

Further reading

Levich, V.G. (1962). Physicochemical Hydrodynamics. Englewood Cliffs, NJ: Prentice Hall. M. Malmsten (ed.), Biopolymers at Interfaces. New York: Dekker (1998). M. Malmsten (ed.), Biopolymers at Interfaces, 2nd Edn. New York: Dekker (2003). Ramsden, J.J. Adsorption Kinetics of Proteins. In: Encyclopaedia of Surface and Colloid Science, A. Hubbard (ed). New York: Dekker (2002) pp. 240–261.

Chapter 2 Chemical and physical attributes of biomedical surfaces For engineering materials, the properties that are important in materials selection for general manufactured objects are well established—density, elastic moduli, thermal conductivity, wear and corrosion rates, etc. Materials selection charts based on these properties are familiar. No such systematic body of knowledge appears to exist in the domain of biomedical materials. The difficulties arise because of the interface of the familiar, relatively predictable, inanimate world with the unfamiliar (in a quantitative, engineering sense) animate world. Indeed, we are all too familiar with the fickleness and unpredictability of thoughts and feelings, and even at the more concrete level of physiology and molecular biology, whose subordination to the laws of physics and chemistry is not usually doubted, the sheer complexity of the systems under consideration often provoke bafflement. Hence there appears to be a quite widely-held notion that biocompatibility is not quantifiable. Clearly such notions, rooted in the reasons sketched out above, do not work to promote a quantitative approach to the selection of biomedical materials. In the past, the “muddling along” approach that is an inevitable consequence of the rejection of the possibility of quantification produced results that were sufficiently acceptable to enable a large and profitable industry of biomedical devices to be built up. Nevertheless, continuing advances in the invention of more and more varied and sophisticated materials, especially thanks to nanotechnology, will increase the visibility of the limitations of the traditional approach. Furthermore, users of devices are becoming more and more critical of their performance, and alternative approaches to health maintenance will mean that providers of devices will have to be increasingly 31

32CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC competitive to maintain their position. It follows that the time is ripe for the introduction of “evidence-based materials selection”, just as “evidence-based medicine” is increasingly demanded when considering whether to approve new drugs for sale. The salient attributes of biomaterials important in their selection are rather different from those applicable to all ordinary engineering materials. Of course, for prostheses the familiar mechanical and thermal properties are also important, but since this book is about surfaces, we do not need to give them attention. Engineering surface properties that are also important for biomedical applications include wear and corrosion rates, and unique to biomedical applications is biocompatibility. The premiss of this book is that biocompatibility is indeed quantifiable (indeed it could have been subtitled “The quantification of biocompatibility”), and it depends on the energetics, morphology and chemistry of the biomedical surfaces whose biocompatibility is under consideration. One of the arguments advanced in favour of the non-quantifiability of biocompatibility has been that it is too dependent on particular environmental circumstances to be a universal parameter, whereas engineering material properties such as density and the elastic moduli are supposed to be intrinsic (at least for terrestrial environments). Nevertheless, at least one significant engineering materials property is environmentally dependent: corrosion rate. Therefore, the necessity of expressing protein adsorption rate (a candidate for quantifying biocompatibility) in terms of the particular protein under consideration should be no more of a hindrance then the necessity of expressing corrosion rate with respect to a particular liquid medium.

2.1

Interaction energetics

In the previous chapter the free energy of interaction between a particle (protein) and a surface on which it might adsorb was discussed, without inquiring more closely into the nature of the interaction. In this section the matter will be examined in more detail.

2.1.1

Pairwise interactions

In the nanorange, in aqueous systems (the only ones we need to consider in connexion with biomedical materials) matter interacts with three different forces: Lifshitz-van der Waals (LW), electron donor-acceptor (da), and electrostatic (el). All are fundamentally electronic in origin. The Lifshitz-van der Waals potential arises through correlations between

2.1. INTERACTION ENERGETICS

33

the electrons of the atoms of adjacent materials. It is therefore generally repulsive, and is weak and short-range, decaying as 1/z, where z is the separation between the two materials. Denoting ∆G(LW)k as the interaction between two infinite planes, the interaction between an infinite plane (representing the surface of the biomedical devices, labelled with a subscript 1) and the sphere of radius r (representing the protein molecule, labelled with subscript 3) is (e.g. van Oss, 1996): (LW)

(LW)k

∆G123 (z) = 2π`20 ∆G123 r/z ,

(2.1)

where `0 is the so-called distance of closest approach, equal to about 1.5 ˚ Angstr¨om units, beyond which the Born repulsion between electron clouds on the atoms of the two materials prevents approach. The electron donor-acceptor force is due to the interaction between Lewis acids and bases in the adjacent materials. Its most common manifestation in biological systems is the hydrogen bond. Electron donors in one material will repel electron donors in the opposing material, but attract electron acceptors. Hence the interaction potential can be both attractive and repulsive. It is strong and long-range, and decays exponentially according to (e.g. van Oss, 1996): (da) (da)k ∆G123 = 2πχ∆G123 exp[(`0 − z)/χ]r , (2.2) where χ is the characteristic decay length for the da interactions. The electrostatic force is the Coulomb force between electrostatic charges created at the surfaces of the materials typically by the ionization of certain elements, or the protonation of the protonation of certain chemical groups, modulated by the presence of dissolved ions, which is taken into account via the Debye length 1/κ, defined by: κ = e(cion /εε0 kT )1/2

(2.3)

for a 1:1 electrolyte such as KCl. The electrostatic interaction potential is given by (e.g. van Oss, 1996): (el)

∆G123 = 4π0 ψ1 ψ3 ln[1 + exp(−κz)]r ,

(2.4)

where the ψ are the electrostatic surface potentials. These can readily be determined using Healy and White’s (1978) surface ionizable group model. For monoprotic acidic or basic groups present at a density Ns on a surface, we have   ±Γ y ± log10 −1 (2.5) pH − pKa = − 2.303 sinh(y/2)

34CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC (taking plus signs for a (Br¨onsted) basic and minus signs for an acidic surface), where y0 = ψe/(kB T ) is the dimensionless surface potential, and Γ is a dimensionless parameter defined by Γ = Ns κ/(4cion ) For an amphoteric surface with isolectric point pH0 (e.g. a metal oxide) we have δ sinh ξ sinh y/2 = Γ 1 + δ cosh ξ

(2.6)

where δ = 2 × 10−∆pK/2 and ξ = 2.303∆pH − y, where ∆pH = pH0 − pH. According to the surface potential and its sign, the immediate vicinity of the surface will be enriched or depleted in protons, resulting in the surface pH pHs = pH +

y . 2.303

(2.7)

In the presence of moderate to high ionic strength liquids between interacting materials, the electrostatic interaction potential is weak and long-range. It can be either attractive or repulsive. Typically, in calculations the net surface charge is taken (mean field approach). From the above it is apparent that to estimate the interaction potentials between a protein and the surface of a biomedical material, it is necessary to know: 1. The Hamaker constant AH (note that ∆G(LW) = −AH r/(6z), for calculating the LW interaction potential; if this is not available it can of course be calculated from the LW component of the surface tension; 2. The dativit`a (electron donor potential), most conveniently expressed as the surface tension component γ ; 3. The recettivit`a (electron acceptor potential), most conveniently expressed as the surface tension component γ ⊕ ; 4. The mean net electrostatic surface potential, assuming that the same data is already known for the protein and the intervening solvent, which will typically be some kind of aqueous solution containing salts and small osmolytes such as glucose. Dupr´e-like laws are then used to compute the pairwise interaction potentials between infinite planar surfaces required in equations (2.1) and (2.2), viz.: (LW)k

∆G12

(LW) (LW) 1/2 γ2 )

= −2(γ1

(2.8)

and (da)k

∆G12

= −2[(γ2⊕ γ1 )1/2 . + (γ1⊕ γ2 )1/2 ]

(2.9)

2.1. INTERACTION ENERGETICS

35

To recall, in these equations subscript 1 refers to the biomaterial and subscript 2 to the solvent with which it is in contact; subscript 3 is used for the protein, and exactly the same equations, mutatis mutandis, are used to compute ∆G13 and ∆G23 . The final step is to compute the overall interaction between protein and biomaterial in the presence of solvent, namely (total)k

∆G123

(total)k

= ∆G22

(total)k

+ ∆G13

(total)k

− ∆G12

(total)k

− ∆G23

.

(2.10)

Some typical values of the surface tension parameters for relevant biomaterials and blood proteins are given in Table 2.1.1.

2.1.2

Protein surface heterogeneity

Application of (2.10) to serum albumin adsorbing on glass producers are well-known prediction that the protein is strongly repelled. The strong adhesion of serum albumin to glass is an equally well-known experimental result. Consideration of the actual amino acid distribution on the surface of serum albumin, as can be deduced from the atomic coordinates now unavailable in the protein data bank, reveals that it is highly irregular (a first glance suggests that it is close to random). Ad hoc attempts to explain the adsorption of serum albumin (net “gamma minus”, i.e. preponderantly an electron donor) to glass (also net “gamma minus”, hence they should be net repulsion) have focused on “gamma plus” patches on the surface of the albumin molecule that can “punch through” the repulsive screen. This approach has several intriguing consequences, such as that the orientation of the albumin molecules on the surface is highly non-random (assuming that once the barrier has been passed there is no orientational randomization), since there is perhaps only one such gamma plus patch. Techniques of molecular microscopy are perhaps not quite yet advanced enough to ascertain whether this is the case in a directly convincing manner. An alternative approach is based on the observation that the actual distribution of gamma plus and gamma minus residues on the surface of the protein is not actually random, and hence the local preponderance of gamma plus or gamma minus depends on the relevant characteristic length of the adsorption problem, which determines the size of the patch whose net Lewis acidity or basicity need to be determined.1 For albumin on silica, the mean field approach gives a strongly repulsive interaction potential; but if a characteristic length of about 1 nm is taken, the potential is close to zero. 1

Calonder, C., Talbot, J. and Ramsden, J.J. (2001). Mapping the electron donor/acceptor potentials on protein surfaces, J. Phys. Chem. B, 105, 725–729.

36CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC

γ (LW)

γ⊕

γ

AH /J×10−20

serum albumin

27

6.3

51

1.28

fibrinogen

37

0.1

38

egg lysozyme

32

4.5

56

29

2.7

60

nylon 6,6

36

0.02

22

polyethylene

33

0

0

polyethylene oxide

43

0

64

polystyrene

42

0

1.1

teflon

18

0

0

cellulose

44

1.6

17

dextran T-150

42

0

55

SiO2

39

0.8

41

TiO2

42

0.6

46

substance Proteinsa

Lipid bilayer lecithin Synthetic polymers

6.58

Polysaccharides

Metal oxidesb

Table 2.1: Surface tensions/(mJ m−2 ) and Hamaker constants of of a selection of substances. Surface tension data mostly from van Oss (1996); Hamaker constants from Dahneke (1972), van Oss et al. (1988) and Roth et al. (1996). a

Hydrated, at pH 7, measured by depositing thick layers on a smooth substrate. b Note that materials like Si, Ti, Al etc. are covered with a thin but tenacious oxide layer under normal terrestrial conditions. Hence these data are practically what is required for calculations relevant to objects made from these metals.

2.1. INTERACTION ENERGETICS

37

A decisive refutation of the adequacy of the mean field approach to calculate on protein-surface interaction given by the classic experiment of Ramsden et al.,2 in which the adsorption and desorption kinetics of two mutants of cytochrome b5, differing only in the position of a single mutated surface charge, not causing any detectable modifications to the structure, were compared at high resolution. The two mutants differed quite significantly from one another, the degree of difference depending on the nature of the adsorbing surface.

2.1.3

Three-body interactions

The traditional concept of complementarity, whether E. Fischer’s original idea of lock-and-key shape complementarity, or the complementarity of positive and negative charges, or that of Lewis acids and bases, or simply the notion that apolar (hydrophobic) patches on proteins would adhere to corresponding apolar patches on the binding partner, whether another protein, another type of macromolecule, or a solid surface, was recently confounded by the discovery, enabled by the amazing progress in X-ray crystallography that has allowed larger and more and more elaborate protein structures to be elucidated, that in certain protein-protein complexes the contacts were between polar patches on one protein and apolar patches on the other. This discovery, which could not be explained according to the existing complementarity paradigm, triggered intensive new reflexion regarding protein interactions. It is possibly the rather pernicious influence of the extensively pursued calculations of protein structures in vacuum—simply because to do otherwise would have completely surpassed the computational capabilities of the time—that there was a regrettable general tendency to overlook the importance of that highly reactive fluid, water, with which proteins are invariably surrounded in their native existence. Thanks to the pioneering work of Max Perutz and others, it is known that the main structural unit of proteins is the alpha helix. Simple polyamino acids such as polyalanine will fold up spontaneously into a single alpha helix. Many proteins, such as the paradigmatical myoglobin, can be adequately described as a bundle of alpha helical cylinders joined together by short connecting loops.3 The alpha helix is held together by hydrogen bonds between the ith and i + 4th amino acid backbones. While denatured myoglobin has no difficulty in refolding itself spontaneously in water, correctly reforming the alpha helices, one might 2

Ramsden, J.J., Roush, D.J., Gill, D.S., Kurrat, R.G. and Willson R.C. Protein adsorption kinetics drastically altered by repositioning a single charge. J. Am. chem. Soc. 117 (1995) 8511–8516. 3 Some proteins have that other principal secondary structure motif, the beta sheet.

38CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC wonder why the process is so robust, when the denatured polypeptide chain is surrounded by water molecules, each of which is able to donate and accept two hydrogen bonds. Although refolding is an intramolecular process, the overwhelming molar excess of water would, one feels, predominate, ensuring that the backbone hydrogen bonds are always fully solvated by the surrounding water. At the same time it is well-known that only an infinitesimal fraction of all possible amino acid sequences are actually able to fold up into a compact three-dimensional structure. Therefore there is something very special amount of sequences that are actually able to fold. One of the intriguing features of natural amino acid sequences is their blend of amino acids with polar and apolar residues. It was formerly believed that the protein folding problem involved trying to pack as many of the apolar residues into the protein interior is possible, in order to minimize the unfavourable free energy of interaction between water and the apolar residue. Of course, referring back to our early discussion of hydrophobicity, the driving force is not that the apolar residues repel water, simply that the cohesive energy of the water is stronger than the relatively weak (van der Waals) apolar-water interaction. At the same time, there are also a significant number (of the order of 50% of the total number of residues) of apolar residues on the surface of a folded globular protein. The presence of an apolar residue in the vicinity of a hydrogen bond is a highly effective way of desolvating it. Folding success actually involves the juxtaposition of appropriate apolar residues with backbone hydrogen bonds. The effectiveness of desolvation of a hydrogen bond can be computed by simply counting the number of apolar residues within a sphere of about 7 ˚ Aradius centred midway between the hydrogen bond donor and the hydrogen bond acceptor.4 This approach, which can be carried out automatically from the atomic coordinates in the protein data bank (PDB), has also revealed the presence of dehydrons, under-desolvated (or “underwrapped”) hydrogen bonds in certain proteins, especially those prone to aggregation, such as the prion protein, sickle cell haemoglobin, cancer-related (i.e. dysfunctional) transcription factors, etc. if a hydrogen bond is surrounded by insufficient apolar residues to effectively desolvate it, it is then vulnerable to binding to whatever collection of apolar groups can effectively carry out the desolvation. Clusters of dehydrons are especially effective “sticky patches”, and the creation of such patches by amino acid mutations is strongly associated with various human pathologies. It is perhaps worth mentioning that dehydron4

Fern´ andez, A. and Scott, R. (2003). Dehydron: a structurally encoded signal for protein interaction, Biophys. J., 85, 1914–1928.

2.2. WETTABILITY

39

related dysfunctions are characteristic of the highest known level of biological organization. The more complex a species, the more components (i.e. different proteins and other macromolecules) it comprises, and the more intricate its web of interactions. As in any multi-objective optimization problem, not all constraints can be adequately satisfied, and the existence of dehydrons or near-dehydrons is one of the penalties of biological complexity. The discovery of the dehydron by A. Fern´andez has thrown into stark relief the limitations of the mean field approach (as epitomized by equation 2.10), which can have only limited utility in predicting protein adhesion (an important element of biocompatibility). The corollary of this more sophisticated view of protein interactions is that artificial surfaces too may not be adequately characterized by mean-field surface energies. The ambiguity associated with the mean field approach is at its lowest for compositionally pure substances, and possibly true solid solutions of multicomponent substances, but it is very likely that composites, especially nanocomposites whose heterogeneity is on the scale of protein surface heterogeneity, must be scrutinized with explicit regard of that heterogeneity.

2.2

Wettability

The content of surface tension, crucial for estimating interfacial energies, is intimately associated with wetting. Wetting means the spreading of a liquid over a solid surface; dewetting is its converse, the withdrawal of liquid from a surface. These are basic processes in countless natural and industrial processes. Wettability is the physical manifestation of interfacial energetics. Although pioneering work in characterizing the interfacial tensions upon which wetting depends was reported two hundred years ago by Young, the processes are still poorly understood. Few experimental techniques are available for investigating the important solid/liquid interfaces: the contact angle method is simple and probably still the most important, but only a handful of laboratories in the world have shown themselves capable of usefully exploiting it. The history of dewetting, a phenomenon of no less industrial importance than wetting, is recent: quantitative experimental work dates from the early 1990’s. An interface may be characterized by a surface or interfacial tension, i.e. free energy per unit area, usually denoted γ. It is essentially intuitive to expect that the spreading of a liquid L on a solid S in the presence of vapour V depends on γSV . The quantitative relationship was given by Young in 1805: γLV cos θ = γSV − γSL . (2.11)

40CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC Fig. 2.1 shows a small droplet in equilibrium on a horizontal surface. The degree of wetting (the term wetting is used regardless of the nature of the liquid; hence dry means not that water is absent, but that all liquid is absent) is inversely proportional to the contact angle θ; θ = 0 corresponds to complete wetting. Young’s equation (2.11) can be easily derived by noting that the surface tension can be written as a force per unit distance. The interfacial forces acting on the triple line T, where three phases S, L, V (solid, liquid, vapour) meet must sum to zero in a given direction (x, parallel to the interface). More formally, it follows from the condition that (at equilibrium) the energies must be invariant with respect to small shifts dx of the position of T. Young Figure 2.1: Small droplets of a liquid L sitting at equilibrium on a horizontal surface S in the presence of their vapour V. (a, lefthand side) on a relatively hydrophobic surface; (b, right-hand side) a relatively hydrophilic surface (assuming L = H2 O).

The structure of T may be very complex. For example, for water containing dissolved electrolyte, the local ion composition may differ from that in the bulk; soft solids may be deformed in the vicinity of T. Nevertheless, Young’s equation, which ignores these details, holds. Length scales for defining θ. The region in which deviations from “far field” quantities occur is known as the core, with radius rC . No simple prescription for determining rC exists, but it is typically ∼ 10 nm. Hence for typical drops used in contact angle determinations with a radius R ∼ 1 mm, the curvature of T may be neglected. Conversely, atomic scale heterogeneity on the ˚ Angstr¨om scale may also be neglected.

2.2.1

The wetting transition

Complete wetting is characterized by a contact angle θ = 0 which implies that γLV = γSV − γSL . (2.12) This is at equilibrium; out of equilibrium, the relation may not hold. A spreading coefficient S may be defined as: S = γSV ˜ − γSL − γLV

(2.13)

2.2. WETTABILITY

41

where γSV is the interfacial tension of a dry solid. Three r´egimes can be ˜ defined: 1. S > 0. This corresponds to γSV > γSV , i.e. the wetted surface has ˜ a lower energy than the unwetted one. (It can be shown that S > 0 is equivalent to the free energy of the SL interface exceeding the LL cohesive energy [6].) Hence wetting takes place spontaneously. The thickness h of the film is greater than monomolecular if S  γLV . The 2 difference γSV ˜ − γSV can be as much as 300 mJ/m for water on metal oxides. Such systems therefore show enormous hysteresis between advancing and receding contact angles. Other sources of hysteresis include chemical and morphological inhomogeneity (contamination and roughness). 2. S = 0. Occurs if γSV ˜ practically equals γSV , as is typically the case for organic liquids on molecular solids. 3. S < 0. Partial wetting. Films thinner than a certain critical value, usually ∼ 1 mm, break up spontaneously into droplets.

2.2.2

Wetting dynamics

Using forced flow in a capillary with velocities v varying over five orders of magnitude, Hoffman established that the capillary number w (equal to vη/γLV , where η is viscosity) depends on θ. At low w, w ∼ θ3

(2.14)

and as w → 1, θ → π. Systematic data using pure nonpolymeric liquids are still lacking. Similar results have been obtained for the spreading of droplets. It has been recognized that the spreading droplet is preceded by a thin film, typically 10–100 nm thick, called the precursor film. It has been detected electrically using conducting liquids.

2.2.3

Dewetting dynamics

Quantitative work dates from 1991, using a smooth fluorocarbon surface on which silicone oils (whose viscosity can be varied over a wide range) or hydrocarbon oils (whose contact angles vary over a wide range) spread by rotation as the experimental system. A hole was then created in the centre of the spread film by aspiration. Most of the important results were obtained by cine photography of the subsequent expansion of the hole (radius rh ).

42CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC Three r´egimes have been identified: viscous, in which the energy released by dewetting (since γSV > γSV ) is dissipated as friction within the liquid, ˜ and hence ultimately as heat; inertial (fast dewetting), in which this energy is converted into liquid kinetic energy (seen for low viscosity fluids such as water, in which the fluid velocity can attain values as high as 100 m/s); and pasty, seen for thick viscous pastes below a certain thickness, which move en bloc as a solid, dissipating the released energy at the interface itself. The phenomenology of viscous dewetting can be summarized in three laws: 1. rh ∝ t, i.e. drh /dt = vh = const. 2. The hole is circumscribed by a bourrelet (i.e. a zone where the film is thicker than elsewhere) which collects the dewetted liquid. 3. vh ∼ θ3 . In the case of inertial dewetting, we have vh = [S/(ρh)]1/2

(2.15)

where ρ is the density of the liquid. For the viscous paste, rh ∝ t2/3 .

2.3

(2.16)

Morphology

Morphology, or shape, can be divided into surface topography and porosity. Ideal surfaces are atomically smooth and flat. In reality, there are likely to be complicated shapes such as peaks and furrows at many different length scales, that depend initially on how the object under scrutiny was fabricated, but which may evolve during service. Undulations at length scales greater than about 10 µm are generally not called rough, but rather “surface waviness”. We shall not be further concerned with these large undulations in this book, since we may presume that they are sensed as planar by most cells, which are themselves smaller than 10 µm. Roughness refers to the unevenness at smaller length scales.

2.3.1

Roughness

At first sight roughness appears to be rather random and may be treated statistically. Although profilometry likely at the continuous description of

2.3. MORPHOLOGY

43

the height h(x) of a surface scanned in the x direction (parallel to the plane of the surface), in this digital age we should assume that it is measured discretely and has a value hi for the ith measuring patch. The quantification of roughness then amounts to a straightforward statistical analysis of the set of height data as summarized in Table 2.2. Symbol

Rq

Definition P mean height (1/N ) N i hi PN arithmetical mean roughness (1/N ) i | hi − hhi | qP N 2 root mean square roughness i (hi − hhi) /N

Rt

maximum peak to valley

max hi − min hi

Rz

mean peak to valley height

Rt averaged over segments

hhi Ra

Name

Synonyms, notes

zeroth moment, D.C. compon

first moment, centreline avera

second moment, interface wid maximum roughness depth

Table 2.2: Table of common roughness parameters. hi is the height of the ith spot on the surface, with reference to some arbitrary fixed level; N is the number of measured spots. Additional parameters are the skew or skewedness (the third moment) and the kurtosis (the fourth moment). More sophisticated surface characterization can be accomplished by determining the power spectral density (PSD) S of surface features, which encapsulates information about the lateral distribution of the individual roughness components: 2 Z L Z L 1 −i2π(fx x+fy y) (2.17) h(x, y)e dxdy S(fx , fy ) = lim 2 L 0 0 where the f are the spatial frequencies in the (x, y) plane of the two dimensional surface profile h(x, y) defined for a square of side L. The power spectral density is the Fourier transform of the autocorrelation function. If there is some evidence of statistical self-similarity, the fractal dimension may be a compact way of representing the roughness. If the power spectral density is plotted (as its logarithm) against the logarithm of the spatial frequency, and if there is a straight line portion extending over two or three orders of magnitude of frequency, then a fractal dimension dC can be derived from the slope. A perfectly flat surface has a (Euclidean) dimension of 2; a fractal surface would have a dimension > 2. The formal definition of fractality relates to the capacity of the surface, and in particular how that capacity the scales with the size of objects placed on the surface. One can determine it by measuring the number n of balls required to cover the surface, using

44CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC a series of differently-sized balls, and plotting log n against log r, where r is the radius of a ball. For a perfectly flat surface, n will be proportional to 1/r2 , but for the fractal surface one will be able to place proportionally more smaller balls, and the slope −dC of the log-log plot but the greater than 2. Experimentally, the fractal surface dimension could be determined by measuring the quantity of nanoparticles of a range of sizes ranging from a few nanometres to a few micrometres required to jam the surface. A fractal dimension is actually the capacity was, as originally defined by Kolmogorov:5 dC = lim ε → 0 log N (ε)/ log(1/ε) (2.18) where N (ε) is the minimum number of p-dimensional cubes of side ε needed to cover the set; p is the dimension of Euclidean space and the set in question is a bounded subset of that space. The fractal concept highlights the importance of characteristic length as a parameter of surface roughness, even though very few real surfaces have self-similarity extending over a sufficiently wide range of lengths to warrant being described as fractals. The fact is that in many of the complex environments with respect to which biocompatibility is being determined, a single roughness parameter is quite inadequate to characterize surface topography. That is perhaps the main reason why the present state of knowledge regarding correlations between surface roughness and cell response is still in a fairly primitive state. Lacunarity An extension of the fractal concept, lacunarity, attempts to capture information about the spatial correlations of real surfaces. The effect of curvature on energetics A corollary of a rough surface is that there will be patches with highly negative or positive curvature. From the interfacial potential equations presented earlier (??) it is clear that increase of curvature will increase the interfacial energy compared with the perfectly flat surface of infinite curvature. Due to the fairly short characteristic lengths of the various types of interaction (Lifshitz-van der Waals, donor-acceptor, electrostatic) this effect only becomes significant at the nanoscale. It is currently a moot point whether the surface tensions derived from contact angles measured on layered films of proteins already incorporate the effective of the small radius of curvature, since 5

A.N. Kolmogorov, A new invariant for transitive dynamical systems, dokl. Akad. Nauk SSSR 119 (1958) 861–864

2.3. MORPHOLOGY

45

one does not really know the molecular structure of these layers. Regarding the other surfaces, of course if the contact angles of the actual surface under scrutiny has been measured, one may assume that the effect of curvature is incorporated into the surface tensions derived from the contact angles, that otherwise the mean curvature of the surface (e.g. the average wavelength of the Fourier description) can be used to correct data obtained from contact angles measured on smooth planar surfaces.

Porosity Porosity is formally defined as the ratio θ of the void volume to the total volume of a given sample of material. It is customary to classify pore sizes as micropores (< 2 nm; the term predates the era of nanotechnology, and hence seems a little inappropriate nowadays: it would be better to call them nanopores); mesopores (between 2 and 50 nm); and macropores (> 50 nm). Open and closed pores are also distinguished: open ones are connected to the exterior of the sample, and closed ones are not. For some practical applications (e.g. porous materials as drug reservoirs, or as adsorbents), only the open pores are relevant. For other applications (e.g. low mass density or low optical density (low refractive index) materials, it is immaterial whether the pores are open or closed. Experimentally it is more difficult to characterize the closed porosity. Porosity is the three-dimensional equivalent of surface roughness. Following on the argument of the previous subsection, one may characterize porosity via a fractal dimension, although the available experimental data may not permit this to be done with acceptable reliability. Porosity, along with roughness, is believed to be important for the assimilation of artificial implants with bony tissue, although the critical parameters for successful assimilation are not presently known. The tortuosity of interconnexions between the porous particularly important for applications in drug-release. In uniform Euclidean systems of any dimension the mean-square displacement hR2 (t)i of a random walker is proportional to the duration of the walk. In disordered systems, on the other hand, hR2 (t)i ∼ t2/dw

(2.19)

with the exponent dw > 2. Hence diffusion in slower than in uniform systems, due to the dead ends, bottlenecks and backbends existing in the disordered structure. The exponent is related to the fractal dimension of the lattice on which the walker is moving.

46CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC

2.4 2.4.1

Chemistry The principles of molecular recognition

The classical lock and key mechanism of biochemical binding specificity first enunciated by Fischer in 1894 appears to have been meant to have been interpreted fairly literally, i.e. a ligand would only bind or “dock” to a receptor if its shape was complementary. The idea of docking governed by shape is of course pervasive in engineering and became particularly widely appreciated through extraterrestrial manoeuvres between spacecraft and space stations. The idea of shape-controlled docking was also offered as the rationalization for the binding specificity of molecularly imprinted polymers. On the other hand, the idea of chemical complementarity was associated from the start with the specificity of nucleic acid binding (base pairing), first enunciated by Watson and Crick. One base presents three (or two) hydrogen bond donors, and the complementary base, which is bound with far higher specificity than any others, presents three (or two) matching hydrogen bond acceptors. Shape is not considered significant. Conceptually, however, one may still give credit to Fischer for the idea: it is a straightforward generalization of shape complementarity. Other biochemical binding, most notably antibody-antigen binding, appears to rely on both shape and chemical (nearly always hydrogen bond donor/acceptor) complementarity. In aqueous solution, this is of course mediated by water, but since water is both a hydrogen bond donor and an acceptor, complementarity is not necessarily destroyed if a layer of water molecules is inserted between the specific binding pair. Most other examples of biochemical binding involving proteins, such as enzymes binding their socalled substrate (‘substrate’ here being used in the sense of a small molecule that is going to be chemically transformed with the help of the enzyme acting catalytically), or indeed most of the cofactors that bind to proteins in order to modulate their activity, rely on shape and chemical complementarity. Sometimes there is also electrostatic complementarity, if ionized groups (most usually protonated or deprotonated) are present. It is possible to analyse the specific binding process according to pattern recognition theory. Suppose that the receptor surface is characterized by a certain arrangement of a hydrogen bond donors and acceptors. Evidently these donors and acceptors are discrete entities and can be labelled εi , the variable taking all possible values d (donor), a (acceptor), or n (neither). A given receptor ρ corresponds to a certain pattern that can be represented as a particular collection {ε} (which we can label Rρ ). Similarly, a certain ligand would correspond to another arrangement of ε, which we can here denote Lλ .

2.4. CHEMISTRY

47

The measure of overlap, i.e. successive binding, for an N -pixel arrangement is then given by N X ρλ −1 Ω =N Lλ ⊗ R ρ (2.20) i

where ⊗ denotes the complementation operator, defined according to Table 2.3. This very elementary first approach neglects of course the spatial correlations within the planes of the binding surfaces. L

R

Result

d

d

0

a

d

1

n

d

0

d

a

1

a

a

0

n

a

0

d

n

0

a

n

0

n

n

0

Table 2.3: Truth table for the complementation operator ⊗. d denotes a hydrogen bond donor, a a hydrogen bond acceptor, and n neither. This is just a first step. An extended truth table should also include electrostatic charges, and the effect of dehydrons should also be taken into account. Electrostatic charges will be influenced by the presence of other ions in the liquid medium that is presumed to bathe the ligand and the receptor. Here it is assumed that shape plays no role, as in a linear or planar array; a more elaborate calculation could of course take shape into account. The ligand-receptor interaction is more complicated in most biochemical binding reactions by virtue of the phenomenon of “induced fit”. This means that after an initial encounter, the ligand or the receptor or both undergo conformational changes in order to improve the value of Ωρλ . One is accustomed to thinking of proteins are the construction with a small number of relatively stable conformational states, hence induced fit is probably a tractable problem. Most theoretical work on the subject has been devoted to explicit molecular dynamics, molecular mechanics or some other kind of computational modelling of the docking process.

48CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC These principles can be equally well applied to the interaction of a biomacromolecul such as a protein with any kind of nonuniform surface. Since however most biomedical materials used until now have been uniform, and it is not been feasible to prepare and characterize with a given patterns of heterogeneity at the nanoscale until very recently, there are very few results in this area so far. Compositional nonuniformity can be characterized by any of the standard diversity indices, e.g. Simpson’s index λ λ=

N X

πi2 .

(2.21)

i

πi can be interpreted both spatially and compositionally. In the former case, the surface is divided up into N pixels and πi is the proportion of a given component in the ith pixel; in the latter, πi is the proportion of the total volume of the particle in the ith (out of N ) types of material. This can be useful when comparing the graininess of compositionally uniform (in the sense that the average composition does not change as one moves about the particle) particles, but is not very useful for characterizing ‘core-shell’ particles and others with similarly non-uniform structures. Another diversity measure is the Shannon entropy: H=−

B X

pi log2 pi ,

(2.22)

i=1

where pi is the fraction of total number of pixels having the grey value lying in the ith grey level, and B is the total number of grey levels. A good candidate for quantifying arbitrarily heterogeneous patterns is lacunarity, which was originally introduced as a way of quantifying the voids in fractal objects; Gefen et al. defined it as the deviation of a fractal from translational invariance;6 Allain and Cloitre consider it as a scale-dependent measure of heterogeneity, i.e. “texture”, of objects in general.7 The basis for determining the lacunarity is, in the simplest case of an image constructed from binary (black or white, lacunarity to values of 0 and 6

Gefen, Y., Meir, Y. and Aharony, A. Geometric implementation of hypercubic lattices with noninteger dimensionality by use of low lacunarity fractal lattices. Phys. Rev. Lett. 50, 145–148 (1983). 7 Allain, C. and Cloitre, M. Characterizing the lacunarity of random and determninistic fractal sets. Phys. Rev. A 44, 3552–3558 (1991).

2.4. CHEMISTRY

49

1) pixels, the distribution n(s, r) of the numbers of boxes of sider containing s white pixels. The lacunarity Λ(r) is defined as Λ(r) = M2 /M12

(2.23)

where M1 and M2 are the first and second moments of the distribution, 2

M1 (r) =

r X

s(r)n(s, r)/N (r) = s¯(r)

(2.24)

s2 (r)n(s, r)/N (r) = σs2 (r) + s¯2 (r) ,

(2.25)

s=1

and

2

M2 (r) =

r X s=1

where s¯ and σs2 are the mean and variance of the distribution respectively and the total number N (r) = (M − r + 1)2 for square pattern size M of boxes of size r, i.e. a type of variance to mean ratio. The lowest possible value of the lacunarity is 1, corresponding to a transitionally invariant pattern (including the special case Λ(M ) = 1). Two approaches to calculating the lacunarity in practice can be recommended: 1. preprocess images by converting them into binary form: defying a threshold T and for each pixel with a greyscale value Z ≥ T set its binary value equal to 1 (white pixel), otherwise said its value equal to 0 (black pixel). Then calculate the lacunarity using the “gliding box” algorithm: a square box of side r is placed at the origin of the image, the 1s within the box are counted to yield the weight of the box s, the box is moved one pixel along the image and the operation is repeated to produce a distribution n(s, r). The box size is then systematically increased and the whole sequence of operations repeated in order to generate the function Λ(r), and more importantly the ‘lacunarity plot’, i.e. a log-log plot of the function Λ(r). Uncertainties may be evaluated heuristically by comparing the lacunarities of different areas of the same size from the same sample. 2. since the values of the lacunarity typically vary with T , multilevel lacunarity may be appropriate. In this procedure, pixels are binned according to greyscale intensity, the number B of bins being varied from a minimum value of 2 (equivalent to the simple lacunarity described above), to a maximum value equal to the number of greyscales. The pixel values are accordingly converted to an integer lying between 0 and the number of bins minus one. The weight s, of the box, is calculated by summing the corresponding values of the pixels in that box.

50CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC The most important inferences from lacunarity are gained by considering the shape of the lacunarity plot, independently from the absolute values at any given box size for any given threshold or number of grey levels within the working region (defined below). From the mathematical definition of lacunarity (2.23), it can be inferred that lacunarity is a function of: 1. the size r of the gliding box—as it increases, many distributions will have lower relative variances and hence lower lacunarities the highest value of the lacunarity occurs for the smallest box size, i.e. Λ(r = 1) = 1/p, where p is the function of occupied sites (except for exceptionally highly cluttered distributions); 2. the average image density (fraction of the pixels in the image occupied by the 1s)—sparse images will have higher lacunarities than dense images, for the same box size; As the distribution was obtained by defining a threshold value T , the lacunarity of an image invariably depends on the threshold value also, which is a corollary of (b) above. (c) the geometric distribution of pixels within the image—a totally regular array obviously has Λ(r) = 1 for all r > the size of the repeating apttern higher lacunarity represents a higher degree of clumping. Comparison of some one and two dimensional structures generated artificially with their corresponding lacunarity plots have suggested the following generalizations:8 1. concave upward curves falling steeply from the maximum value at r = 1 and rounding off as r → ∞ correspond to random arrangements; 2. convex upward curves correspond to more structured arrangements: the lacunarity gently falls to a shoulder a the value of r corresponding to the characteristic scale of the pattern, and then falls steeply before rounding out to Λ = 1; 3. self-similar (fractal) patterns have long linear or quasilinear stretches on the log-log lacunarity plot with a slope equal to the difference between the fractal and Euclidean dimensions [?]. Although a key feature of lacunarity is its ability to capture texture over many scale lengths, it is nevertheless sometimes useful to reduce it to a single parameter which can be compared with some other variable characteristic of 8

Plotnick, R.E., Gardner, R.H., Hargrove, W.W., Prestegaard, K. and Perlmutter, M. Lacunarity analysis: a general technique for the analysis of spatial patterns. Phys. Rev. E 53, 5461–5468 (1996).

2.4. CHEMISTRY

51

the process creating the pattern. Examples of such semen parameters are: (i) the mean radius of curvature of the lacunarity plots; and (ii) the mean values of the first and second derivatives of the lacunarity plots.

2.4.2

Corrosion

Corrosion is essentially an electrochemical phenomenon that applies to metals. It refers to the anodic dissolution (oxidation) of the metal, releasing electrons. For corrosion to take place, a sink of electrons (cathodic reaction) needs to be coupled to the anodic reaction. A common cathodic reaction is the reduction of dissolved oxygen. Hence ensuring the absence of oxygen is often an effective way of preventing corrosion from taking place. The electrochemical series (Table 2.4) orders the metallic elements according to their propensity to release electrons. Those with very high potentials are naturally resistant to corrosion. It is of course well known that the most corrosion-resistant metals are either too expensive or mechanically too weak to be useful in biomedical applications. One other hand, some of the most important metals such as titanium are apparently highly prone to corrosion. For acceptable corrosion resistance it is important that such metals form thin protective passive oxide films that act as a barrier separating the metal from its environment. The metals and alloys often used for biomedical applications, such as 316L stainless steel, titanium, Ti-6Al-4V, Co-Cr-Mo and Ni-Ti, do indeed form passive surface oxide films within the human body, which usually provide good corrosion resistance. Problems can arise if the oxide film is damaged, possibly by mechanical abrasion. In general the film will reform, but the is only useful if it happens faster than corrosion. Once there is damage, corrosion can be quite rapid because small damaged area is likely to have a high current density of electrons, balanced by a low current density of electrons over the remaining undamaged surface. Small imperfections are therefore quite likely to lead to pitting. If the damaged area does not repassivate, fretting corrosion is said to occur. Galvanic corrosion takes place when metals of different electrode potentials are placed within the same local zone. The Pourbaix diagram can be used to predict that stainless steels are likely to undergo corrosion in many biomedical environments. However, titanium should remain passive in almost all aqueous solutions, a major advantage that accounts for its popularity. Even titanium does degrade, however (see Figure 2.2. As well as causing weakening of the implant, corrosion also has the undesirable consequence of releasing possibly toxic metal ions into the body, at concentrations far higher than in those normally encountered. Data on

52CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC element

E0 vs s.h.e./V

Gold

1.43

Platinum

1.20

Mercury

0.80

Silver

0.79

Copper

0.34

Lead

−0.13

Molybdenum

−0.20

Nickel

−0.25

Cobalt

−0.28

Iron

−0.44

Chromium

−0.73

Zinc

−0.76

Aluminium

−1.33

Titanium

−0.63

Magnesium

−2.03

Table 2.4: The electrochemical series of metals used in biomedical applications. Standard potentials are given versus the hydrogen electrode (standard potential = 0).

the topic is uneven in coverage and of variable quality. A table in Ratner et al.’s Biomaterials9 gives some figures suggesting that significantly elevated concentrations (i.e. about 50 times greater than the normal level found in implant-free humans) of titanium, the most likely material to have been used for the implant (cf. Table 3) are only found in serum in cases of patients with poorly functioning (and requiring surgical revision) total knee arthroplasties. A similar enrichment is found in the synovial fluid.

9

Ratner, B.D., Hoffman, A.S., Schoen, F.J. and Lemons, J.E., 2004, Biomaterials Science: an Introduction to Materials in Medicine, Academic Press; Table 5, p. 547; no information on the material of the giant replacement is given.

2.5. BIODEGRADABILITY

53 mam1-1

Figure 2.2: Wear of the metallic head and ferulene bush of a novel endoprosthesis 10 years after initial implantation (source: Sytenko Institute of Spine and Joint Pathology, Kharkov.

2.5

Biodegradability

It has only relatively recently that the synthetic biodegradable materials have become sufficiently strong and otherwise reliable enough to be considered as useful candidates in biomedical applications. There are strong potential advantages in using biodegradable materials, which ultimately disappear from their host. In nearly every case the body in principle has the capacity to repair itself; the notion of the artificial degradable implant is to provide a scaffold during the transition period of natural reconstruction (during which the body might otherwise be in a life-threateningly weak condition). Such an approach obviates the problematic chronic inflammatory or encapsulatory responses of the body to many permanent implants. Most biodegradable materials are synthetic polymers (although natural polymers, or at least modified natural polymers or polymerized natural monomers, are being increasingly explored) “Biodegradability” can generally mean one of two distinctive processes: actual breaking of bonds by by chemical activity within the host, perhaps even resulting in the fragments being inserted into the regular metabolism of the host; or the simple slow dissolution of the material, which is then ultimately excreted. The material called polylactic acid has attracted the most attention. It was first used several decades ago for sutures and temporary bone fixations. Lactic acid can also be copolymerized with glycolide monomers to extend the chemical diversity. The polymers are insoluble in water, but can be hydrolysed. The actual degradation rate depends on the degree of crystallinity as well as the molecular weight and the actual chemical composition. Typical degradation times range from days to years. One notices that in the table of commonly used materials (Table 3) biodegradable polymers do not appear. Degradation unfortunately introduces some new problems, such as granulation during the aggregation processes, releasing particles that could migrate elsewhere within the body. The search to find degradable materials with acceptable performance is currently an active research area. Even non-degradable polymers may leach unreacted monomer or plasticizer into their environment in a passive, diffusion-driven process. But many

54CHAPTER 2. CHEMICAL AND PHYSICAL ATTRIBUTES OF BIOMEDICAL SURFAC “durable” materials are also attacked by enzymes that can be expressed, either by the body’s own cells, or by symbiotic bacteria (and, in principle, invasive bacteria, but if these reach the concentration sufficient to be effective for degrading artificial implants, the body is also likely to be suffering physiologically from the invasion). The genome plasticity of bacteria in particular can result in rapid optimization of an enzyme able to exploit an implant as a food source, for example. The main enzyme-catalysed degradation processes are hydrolysis and oxidation.

2.6

Further reading

Cacace, M.G., Landau, E.M. and Ramsden, J.J. (1997). The Hofmeister series: salt and solvent effects on interfacial phenomena, Q. Rev. Biophys., 30, 241–278. S. Havlin and D. Ben-Avraham, Diffusion in disordered media, Advances in Physics, 36 (1980) 695–798. Healy, T.W. & White, L.R. (1978). Ionizable surface group models of aqueous interfaces. Adv. Colloid Interface Sci. 9, 303–345. T.A. Loomis and A.W. Hynes, Essentials of Toxicology, 4th edn, San Diego: Academic Press (1996). van Oss, C.J. (1996). Forces interfaciales en milieux aqueux. Paris: Masson. van Oss, C.J., Chaudhury, M.K. and Good, R.J. (1988). Interfacial Lifshitz-van der Waals and polar interactions in macroscopic systems. Chem. Rev. 88, 927–941. Ramsden, J.J. (2000). The specificity of biomolecular particle adhesion. Colloids Surf. A 173, 237–249. P. Schaaf, J.-C. Vogel and B. Senger, Irreversible deposition/adsorption processes on solid surfaces, Ann. Physique 23 (1998) 1–89.

Chapter 3 Biofluid compatibility This chapter is concerned with processes that occurr when biomaterials are in contact with biofluids. Many artificial devices introduced intio the human body spend their time in contact with one of the various liquids that circulate in the body. The most important of these is the blood, an extremely complex fluid not only because of the huge variety of macromolecules dissolved in it, but also because it contains large numbers of circulating cells of several types. In comparison, urine is much simpler. Finally, although their volume is comparatively tiny, we shall also consider lachrymal secretions, and the buccal environment (saliva), as well as that of the stomach. The main area of application of this knowledge is to in vivo devices such as stents and heart valves past which blood is flowing, and ex vivo devices such as kidney replacement dialysis machines used for treating blood. The problem of biofluid compatibility is therefore a direct application of the principles developed in Chapter 1. The biomedical surface design issues typically concern the suppression of biomacromolecule adsorption, or, in the case of kidney dialysis machines, selective permeability or adsorption or both. Once again, it can be emphasized that biomacromolecular recognition is the fundamental process that needs to be understood. All cells are enveloped by their own surface biomacromolecules, and therefore the interaction of a cell with any artificial surface is primarily determined by the interaction of those cell surface biomacromolecules with the artificial substrate. It should also be mentioned that even though early on in this book a distinction was made between biocompatibility of materials destined to be implanted in tissues and materials destined to remain in contact with the blood, most tissue implants have to be introduced invasively into the body, and therefore unavoidably come into contact with the blood, maybe only for a few minutes, but this is sufficient for them to become coated with blood proteins, so rapid is the process of adsorption. This inevitability poses an 55

56

CHAPTER 3. BIOFLUID COMPATIBILITY

especial problem for biomedical surface designers, because essentially the material surface has to fulfil two opposing functions: resistance to (blood) protein adsorption, and promotion of (tissue) protein adsorption.

3.1

Haemocompatibility

Haemocompatibility, defined as ‘tolerant of blood’, is a particular aspect of biocompatibility. The materials from which medical devices implanted into the bloodstream are made must be haemocompatible in order to be serviceable. In this section, the basic principles outlined in Chapter 1 will be extended and refined in order to make them fully applicable to problems of haemocompatibility.

3.1.1

The composition of human blood

Table 3.1 gives the molecular, ionic and cellular composition of human blood. The blood is transported throughout the body by the vascular system— arteries bearing oxygenated blood from the heart, and veins bearing blood that has discharged its cargo of oxygen in tissues back to the heart. Thus it might be thought that the primary function of the blood is to carry oxygen, which is of course essential to life, and the other functions of transporting nutrients, metabolites, excreta etc. are exaptations. This book is not the place to expand on the discussion of this point. An analogy with a network of roads, railways, rivers or canals carrying merchandise is quite apt: the analogy of the telegraph or telephone network carrying information is then the nervous system. Constant repair and defence work is necessary to maintain the system, just as with the transport networks built by humans. Although the blood is not normally accessible to the external world, any damage to the skin typically results in a wound that opens an access channel between the external world and the blood, through which microbes, dust particles etc. may enter into the vascular network. As will be described in a little more detail below, constantly circulating macrophages will phagocytose and attempt to destroy, usually with success, the invading entity. Another population of cells, the platelets, will swiftly act to repair any breaches of the endothelial lining of the vascular network.

3.1. HAEMOCOMPATIBILITY

entity Proteins (Mr )

57

concentration

function

mg/mL

albumin (66 500)

40–55

general scavenger of smaller molecules

fibrinogen (3.4 × 105 )

2–4

participates in thrombus formation

fibronectin (4.7 × 105 )

0.15–0.3

immunoglobulin G (150 000)

5–15

part of the adaptive immune response

(80 000)

2–3

iron transport

Ions

mM

Ca2+

2.4

Na+

80–90

+

40–60

transferrin

K

Cl− osmolytes glucose

80–85 g/L 0.8–1

urea

0.2–0.4

Cells

µL−1

erythrocytes

food metabolite

4.5 − −6 × 106 3

neutrophils

3 − −6 × 10

lymphocytes

1.5 − −3 × 103

platelets

1.5 − −3.5 × 105

oxygen transport destruction of microbial invaders thrombus formation

Table 3.1: The composition of human blood. Data is for whole blood, except the protein concentrations, which are given for plasma (serum), i.e. blood from which the cells have been removed.

58

CHAPTER 3. BIOFLUID COMPATIBILITY

3.1.2

Initial recognition events

Initial detection of events requiring intervention take place through that signalling mechanism known as molecular recognition (pattern matching). The circulating cells or macromolecules are constantly weakly interacting with their surroundings; occasionally they will encounter a complementary match to their own patterns, and bind with higher affinity. Very often the “receptor” is the transmembrane protein, and the ligandbinding event triggers a conformational change in the protein, resulting in a sequence of reactions in the interior of the cell (a “cascade”—see the next subsection) that ends up with the synthesis of some new1 protein. Once a cell or macromolecule is resident on, say, a patch of the epithelium, binding partners from either the circulating or the stationary entity can be recruited to the binding site. This means that congeners of either the “ligand” or “receptor” (often the two molecules are actually very similar), that are anyway diffusing in the outer membrane of the cell, arrive at the binding site and form a cluster of binding pairs. The subsequent events triggered in the interior of the cell, possibly leading to the synthesis of specific proteins within the cell, as already mentioned, are dependant on the formation of the cluster. This is a convenient way of increasing the ratio of input and output signal-to-noise ratios.2

3.1.3

The coagulation cascade

The function of coagulation is to seal any breaches in the vascular system. As haemophiliacs well know, failure of the coagulation mechanism leads to uncontrolled bleeding from any cut or similar wound, with possibly fatal consequences. The end result of the so-called coagulation cascade is the formation of a solid plug of matter, called a thrombus, that is a composite of cells (platelets) and protein (mostly fibrin, i.e. polymerized fibrinogen). The thrombus is essentially a composite structure made from platelets (“bricks”) and fibrin (“cement”). The crucial polymerization is effected by an enzyme called thrombin. Since fibrinogen is circulating constantly in the blood, it would be fatally dangerous to have thrombin systemically present. Adventitious coagulation forming a clot poses the risk of blocking a vital blood vessel. Therefore, either the thrombin has to be normally absent, and synthesized within an appropriate cell as a result of triggers ending up with a transcrip1

In the sense that it was not present before, or only present at very low concentration. This is a principle of great generality in molecular signalling. See Manghani, S. and Ramsden, J.J., The efficiency of chemical detectors. J. Biol. Phys. Chem. 3 (2003) 11–17. 2

3.1. HAEMOCOMPATIBILITY

59

tion factor binding to a promoter, or constantly present in some precursor form that can be swiftly activated. Clearly either approach has its merits, crucial parameters being swiftness of response and risk of inadvertent triggering, and both mechanisms can be found in biology, that for the particular process of blood coagulation, nature has chosen the latter course. The protein prothrombin is circulating everywhere, but it needs to be converted to the active enzyme thrombin by the action of another enzyme. Clearly that enzyme in turn cannot be systemically present (otherwise prothrombin would constantly be being converted to thrombin), hence that enzyme in turn is normally present in the form of an inactive precursor, that needs to be enzymatically activated. This sequence of reactions constitutes a cascade, and is shown schematically in Figure 3.1. cascade Figure 3.1: Schematic diagram of a molecular cascade.

Given that each reaction is stochastic (and typically in a very real sense, because the actual numbers of reacting molecules in the relevant volume may be very small), it is appropriate to describe the cascade using a set of nested Markov chains (also called hidden Markov model, HMM). Let A represent the inactive state of an enzyme, and A* its activated form, due to the action of an environmental perturbation E whose activity is represented by the variable 0 ≤ e ≤ 1. In the complete absence of E, the enzyme has no chance of becoming activated, but in its presence the probability in unit time of becoming activated is e (some other function gives the relationship between a particular quantity of E in the environment of the blood vessel or cell and e). This can be represented by a two state Markov chain: →

A

A∗

A

1-e

e

A∗ 0.5 0.5 Once the activated form exists, it will have a finite lifetime (here represented by the spontaneous decay probability equal to one half of returning to the ground state in unit time). The enzyme A* is able to catalyse the formation of an active enzyme B* from an inactive circulating precursor B. In the absence of A*, B will remain

60

CHAPTER 3. BIOFLUID COMPATIBILITY

inactive. The matrix of transition probabilities for B will look like

→

B

B∗

B

1-a

a

B∗ 0.5 0.5

and so on for the reaction catalysed by B* et seq. In reality, very often all the transition probabilities depend on other factors, and there may be additional feedbacks (e.g. B* might complex with E, lowering e). This model can be equally well used to describe the intracellular signalling cascades that convey information from the exterior of the cell into the nucleus. In that case, typically the active form of the enzymes are phosphorylated, hence the activating enzymes (which must in turn themselves be phosphorylated) are the so-called kinases. The intracellular cascade requires de-activating processes, which are supplied by the so-called phosphatases. Were they not present, the cell would quickly run out of non-phosphorylated precursors, and moreover arresting the signal could only take place by the relatively slow process of protein degradation. The primordial reason for having cascades is to raise the molecular detection efficiency of the process. Fundamentally, what is happening is that a signal (the presence of a certain entity external to the cell) is being detected, and information embodied in the signal has to be channelled to the nucleus. The exterior of the cell is however very crowded, with many molecules that are all rather similar to one another, and all of them have a weak affinity to the exterior of the cell. Although only the binding of a specific molecule will result in a definite conformational change of the receptor, random thermal motions are happening all the time, and to prevent false alarms, it is essential to impose a signalling threshold. The cascade, in conjunction with the clustering, is a convenient way of achieving this. Noticed that the cascade results in molecular amplification: one activating enzyme can activate many molecules, and such local predominance also ensures that only the correct signals are acted upon. It is exactly analogous to a human signaller trying to transmit an alphabetical message in a very noisy environment: an effective way of achieving fidelity is to repeat every letter say five times, e.g. the word “fidelity” would be transmitted as “fffffiiiiidddddeeeeellllliiiiitttttyyyyy”. There are several cascades that are part of the defence mechanisms within the blood. They are summarized in Table 3.2

3.1. HAEMOCOMPATIBILITY

61

Name

stimulus

result

complement

polysaccharides, −vely charged surfaces membrane attack ⇒ cell lysis

contact activation

polysaccharides, −vely charged surfaces vasodilation

clotting

factor XII adhesion

thrombus formation

Table 3.2: The principle blood cascades.

3.1.4

The immune system

The immune system is an extraordinarily complicated mechanism for neutralizing external attack. “External” here signifies, in the widest sense, anything abnormal with respect to the regular functioning of the body. Thus it could be literally invasion by a microbial pathogen, or a mineral particle, entering the body from the air, or via ingestion of food or drink; it could be the mutation of a member of the native bacterial flora of the body into a pathogenic form; it could be the malfunctioning of one of the body’s own cells (apoptosis, necrosis, tumorigenesis etc.), or it could be bruising of tissue caused by some external impact. There are several general requirements for such mending apparatus, including swiftness of response, but it should also not impose an unbearable burden on the body, and it should not interfere with those parts of the body, normally the overwhelming majority, that a perfectly normal. From an initial viewpoint, it is very clearly and multi-objective optimization problem, and clearly compromise is unnecessary. But one cannot but admire the sophistication with which nature has largely solved the problem of dealing with the constant small attacks that are an inevitable consequence of living in a richly varied ecosystem. Clearly such a system presents a rich and challenging object for cybernetic modellers. It must however be emphasized that any reasonable model must be open-ended. New pathogens are constantly arising due to the extraordinary genomic plasticity of the microbial world, and new chemicals and synthetic particles are constantly being introduced into the natural ecosystem by man’s own industrial activities. Most engineered systems are only able to deal with a fixed, predefined repertoire of eventualities. We may presume that any primitive immune system built on such principles would long ago have been eliminated by natural selection. Many microbial pathogens have the egregious habit of unrestrained proliferation. That is the primary reason why swiftness of response is important. The sooner any invasion can be suppressed, the less work there will be to

62

CHAPTER 3. BIOFLUID COMPATIBILITY

do. This concept is again familiar in many everyday situations. For example, the sooner delay can be made up in the running of a railway train, the fewer problems there will be in having to reschedule paths and deal with disgruntled passengers who have missed connexions. The immune system can be divided into two classes: the “innate” and the “adaptive”.3

3.1.5

The innate immune system

The so-called innate system, because it is found in lineages much more ancient than our own, is often considered to be primitive (compared with the adaptive immune system, which is only found in the highest organisms). “Primitive” however belies the great sophistication of the innate system considered in its own right. The primary constituent of the innate response is the phagocyte, comprising essentially macrophages and neutrophils. As the name suggests, these cells are able to ingest and destroy invading microorganisms. A glimpse of the sophistication of this ‘primitive” response is shown by the fact that concomitantly with the ingestion process they release cytokines, small proteins that bind to receptors on the other circulating cells, resulting in a constantly amplifying and adapting response to the invader. The macrophages themselves typically then start acting as antigen-presenting cells (APC),4 an essential part of the adaptive immune response (see below). The actual process of ingestion and destruction takes place via endocytosis, which internalizes the invader, and inside the macrophage lysosomes fuse with vesicles containing the invader and destroy them by a combination of digestive enzymes and toxic free radicals, typically the so-called reactive oxygen intermediates (ROI).5

3.1.6

The adaptive immune system

The main cellular components of adaptive immunity are: 3

Occasionally, physical and chemical barriers to invasion, such as the skin and the macromolecular constituents of the mucosal surfaces that can engulf and neutralize microbes, are also considered to be part of the innate immune system. Similarly, the cascades described in the previous subsection are also sometimes grouped under heading of the innate immune system. 4 Antigen presentation is the main function of the so-called dendritic cells. 5 If the macrophages are unable to phagocytose a foreign particle, they may fuse together around it to create a syncytium called a ‘foreign body giant cell’, which may have dozens of nuclei. This typically happens in the case of large blue asbestos fibres.

3.1. HAEMOCOMPATIBILITY

63

1. helper T lymphocytes (or helper T cells) that excrete cytokines, small proteins that bind to other cells, modulating their activity; 2. cytotoxic T lymphocytes that can kill selected cells (called the cytotoxic T lymphocyte (CTL) response), e.g. by releasing perforins, proteins that perforate the cell membranes of the enemy target; 3. B lymphocytes, that make antibodies. The main protein components are: 1. the major histocompatibility complex (MHC), a family of proteins found on the surface of cells that form a complex with oligopeptides (“epitopes”) derived from a sample of the internal proteins of the cell; 2. the cytokines secreted by the macrophages and T lymphocytes; 3. the perforins released by killer T cells; 4. the antibodies secreted by the B lymphocytes. A particular T cell will have thousands of copies of a receptor (called T cell receptor, TCR) for a particular peptide-MHC complex. All lymphocytes originate in the bone marrow. The T cells then pass to the thymus (whence the name), where those cells with an affinity for self, i.e. the cells of the body) are eliminated (this process has been called “tolerization” by Forrest and Hofmeyr;6 it can be modelled by supposing that if the immature receptors had an affinity for, and hence bound to, any of the host’s peptide-MHC complexes, that binding could be detected and lead to the elimination of the offending cell). Those that survive the tolerization process pass into the bloodstream. Like any cell, they have a finite lifetime, and if during a lifetime they are never stimulated by finding a peptide-MHC complex with which they can bind, they will simply be allowed to die. On the other hand, suppose that can one of the host cells of the body is infected by a virus and therefore expresses a non-self peptide-MHC complex to which one of the T lymphocytes can bind: the T cells were then replicate and attack the infected cell. This is called the primary response. The responsive cells prolong their longevity by becoming memory cells; any subsequent attack by the same infectious agent can make use of that memory and mount a much more rapid response in the primary one.7 This memory that gradually builds 6

Immunology as information processing, in: Design Principles for Immune System and Other Distributed Autonomous Systems, L.A. Segel and I.R. Cohen (eds), pp. 361–387. Oxford: University Press (2000). 7 This is the principle of vaccination.

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CHAPTER 3. BIOFLUID COMPATIBILITY

up (the “immunome”) throughout our lives is as much part of our individual personalities as our brain-based memory. Each of our immunomes is unique— constituting a record of the foreign attacks to which we have been subjected throughout our lives. Potentially, this could cause great variability in the response of individuals to foreign implants—but mapping the vast immunome of a typical mature individual is an even more daunting task then mapping the proteome. Immunomic variability is likely to be much more significant than genetic variability among humans. Nevertheless, ultimately one should envisage personalized implants (especially those with “smart” features), in the same way that personalized drugs are being increasingly considered. The B lymphocytes meanwhile mature in the bone marrow (the bursa in birds) (whence their name). They produce antibodies f the classes labelled G, M, E etc. These are immunoglobulin (Ig) molecules with general composition HL2n , n = 1, 2, 5, where H denotes the so-called heavy polypeptide chain and L denotes the so-called light chain, which includes the variable amino acid epitopes that bind to the antigens. Each member of the population of B cells produces a different antibody (Ab), and an encounter between the B cell and its antigen leads to clonal expansion of that cell type and massive amplification of antibody production. Antibodies binding to an antigen interact with a relatively small portion of the molecule. The number N of foreign antigens which must be recognized by an organism is very large, perhaps greater than 1016 , and there is a smaller number N 0 (∼ 106 ?) of self-antigens which must not be recognized. Yet the immunoglobulin and T-cell receptors may only contain n ∼ 107 different motifs. Recognition is presumed to be accomplished by a generalized lock and key mechanism involving complementary amino acid sequences. How large should the complementary region be, supposing that the system has evolved to optimize the task? If PS is the probability that a random receptor recognizes a random antigen, the value of its complement PF = 1 − PS maximizing the product of the probabilities that each antigen is recognized by at least one receptor, and that none of the self-antigens is recognized, i.e. 0 (1 − PFn )N PFnN , is:8 −1/n  N . (3.1) PF = 1 + 0 N Using the above estimates for n, N and N 0 , one computes PS ≈ 2 × 10−6 . Suppose that the complementary sequence is composed of m classes of amino acids and that at least c complementary pairs on a sequence of s amino acids 8

Percus J K, Percus O E and Perelson A S Predicting the size of the T-cell receptor and antibody combining region from consideration of efficient self–nonself discrimination 1993 Proc natl Acad Sci USA 90, 1691–1695

3.1. HAEMOCOMPATIBILITY

65

are required for recognition. Since the probability of a long match is very small, to a good approximation the individual contributions to the match can be regarded as being independent. A pair is thus matched with probability 1/m, and mismatched with probability 1 − 1/m. Starting at one end of the sequence, runs of c matches occur with probability m−c , and elsewhere they are preceded by a mismatch and can start at s − c possible sites. Hence PS = [(s − c)(m − 1)/m + 1]/mc .

(3.2)

If s  c > 1, one obtains c = logm [s(m − 1)/m] − logm PS .

(3.3)

Supposing s to be a few tens, m = 3 (positive, negative and neutral residues), and again using the numbers given above (since they all enter as logarithms the exact values are not critical) one estimates c ∼ 15, which seems to be in good agreement with present experimental knowledge. The above has only given a brief glimpse of the amazing complexity of the immune system. Our present knowledge is in any case of imperfect, and new discoveries are constantly being made, despite more than a century of intensive investigation. For example, very recently it has been ascertained that the small interfering RNAs, whose very existence was only discovered a few years ago, also play a role in the immune system.

3.1.7

Haemodialysis to combat renal failure

The concept of the artificial kidney is to pass the blood through an external loop containing semipermeable membranes or adsorbents or both that remove the toxic components constantly accumulating in the blood of the diseased patient whose kidney is no longer able to deal with them.9 Haemodialysis was introduced as a therapy around 1960 in the UK, and is currently used to treat about 40% of patients suffering from established (irreversible) kidney failure, which means less than 10% of normal function. The only alternative treatment is transplantation of a donor kidney, but there is a constant shortage of donors. Compared with the sophistication of the natural kidney,10 The action of the dialyser is very crude. Its designed main function is to remove urea. Essential vitamins and amino acids are also removed; on the other hand the so-called middle molecular weight (500 < mR < 50 000) molecules are 9

C. Jacobs, C.M. Kjellstrand, K.M. Koch and J.F. Winchester (eds), Replacement of Renal Function by Dialysis, 4th edn. Dordrecht: Kluwer (1996). 10 S.R. Thomas, Modelling and simulation of the kidney, J. Biol. Phys. Chem. 5 (2005) 70–83.

66

CHAPTER 3. BIOFLUID COMPATIBILITY

removed only partially and indiscriminately. In the context of this book, the main problem is fouling of the dialysis system, not only the membrane itself but also the associated tubing. This problem will continue to the present even if in dialysis membranes are replaced by selective adsorbents.

3.1.8

Stents

These devices (see Figure 3.2) are used to support weakened arterial walls and prevent them from collapsing, to open up cardiovascular channels that have become coated with residues such as plaque, or to give temporary relief of urinary tract obstruction. They are inserted by the surgeon into the blood channel all urinary tract in collapsed form, and then expanded. The two main materials in use are NiTiNOL and stainless steel. They are handled in quite different ways. stent Figure 3.2: Examples of modern stainless steel stents. NiTiNOL (an acronym for NIckel TItanium Naval Ordnance Laboratory) is the name of a family of alloys containing a nearly equal mixture of nickel (typically 55% w/w) and titanium, and possibly traces of other elements. It possesses the remarkable property of “shape memory”, i.e. it can be formed in a certain shape (in this application, the shape of the stent in its final form), plastically deformed cold (e.g. in iced water), and then where it is heated (at body temperature), its original shape is restored.11 The stainless steel stents are inserted in collapsed form, with a deflated balloon in the middle. Once the stent is in place, the balloon is expanded, forcing the stent to expand in turn, and the balloon is then deflated and withdrawn. This procedure places a particular requirement on the stent to be completely burr-free.

3.2

Urine compatibility

The composition of healthy urine is much simpler than that of the blood. Its main constituents are given in Table 3.3. 11

The memory is is due to a crystalline phase change known as a thermoelastic martensitic transformation. Below the transformation temperature, NiTiNOL is martensitic and soft (its microstructure is characterized by “self-accomodating twins”, a zigzag like arrangement of microcrystals). Martensite is easily deformed by de-twinning. Heating the material converts the material to its high strength, austenitic condition.

3.3. LACHRYMAL SECRETIONS

67

In the main implants inserted into the urinary tract are catheters and prostate stents. The latter are preferably made from NiTiNOL shape-memory alloy, also popular for stents in blood vessels. They are typically indicated for cases of benign prostatic hyperplasia (enlarged prostate), which tends to obstruct bladder function. Trials with biodegradable stents (dissolvable polyglycolic acid) have been disappointing for a variety of reasons. Catheters represent a far more numerous application. They are typically used in connexion with incontinence. Their main mechanical requirement is flexibility. This dictates that they are usually made of some kind of rubbery polymer. Clinical experience has shown that catheterization very commonly leads to urinary tract infection—in many hospitals this is indeed the most common nocosomial infection. The main source of infection is colonization by the bacterium Proteus mirabilis. They colonize the surface of the catheter, forming the usual type of bacterial biofilm in which the bacteria are embedded in a polysaccharide matrix. P. mirabilis excudes urease, which generates ammonia from the urea abundant in urine, raising the pH and creating alkaline conditions. This causes the precipitation of minerals, especially magnesium ammonium phosphate (struvite) and calcium phosphate (hydroxyapatite). Crystals of these minerals become incorporated in the organic matrix of the biofilm to form a growing and durable incrustation that eventually blocks the catheter. All materials typically used for catheters, i.e. mainly silicone, possibly coated with Teflon or a hydrogel, appear to be susceptible to incrustation. This problem has currently not been solved. Uropathogenic strains of Escherichia coli are also known to colonize the bladder and its associated tracts, often forming robust biofilms in which the bacteria are encapsulated in a polysaccharide-rich matrix in turn surrounded by a protective shell made from uroplakin. Because of the lack of urease formation however, this bacterium is not associated with the biomineralized incrustation so characteristic of P. mirabilis infections.

3.3

Lachrymal secretions

The main implant of relevance to the eyes is the contact lens, which is currently used in vast numbers worldwide. The main physical attribute required is perfect optical transparency. This sharply limits the range of materials that can be used. Since contact lenses are typically replaced daily, or at least taken out of the eye and placed in a cleaning solution, long-term stability and resistance to biopolymer adsorption is not an issue.

68

CHAPTER 3. BIOFLUID COMPATIBILITY

3.4

Buccal secretions

This section is concerned essentially with dental implants. Although the contents of the mouth (Table ??) rarely have the complexity associated with the thousands of different macromolecules found in the blood, on the other hand the mouth is, more than any other of the fluid-containing cavities of the body, open to the external world, and is therefore unregulated, except by the volition of its owner.

3.5

Bioanalytical surfaces

The growing prevalence of largely automated analytical devices for clinical analysis, replacing previous manual procedures, including high throughput devices using only very small sample volumes and miniature biosensors using even smaller ones, have led to the development of a new branch of biomedical surface science and technology that is wholly concerned with clinical samples handled in vitro. There are two main classes of instruments: 1. those exemplified by the Roche “COBAS” device: the clinical samples (e.g. blood, urine) are preloaded in large numbers onto the device, which mixes them with reagents and measures the optical absorption or fluorescence, statically or kinetically, to determine the concentrations of the analytes of interest; 2. biosensors for point-of-care use, in which the sample is typically analysed via interaction with a bioactive surface (e.g. the glucose biosensor is coated with the enzyme glucose oxidase, with which the glucose in the blood reacts). In either case, the device surface to sample volume ratio is high. Inappropriate surfaces may, in the worst case, cause of the entire loss of the sample of the analyte of interest before it can be analysed. Even partial adsorption of analyte may cause severe distortions in the outputs of the instrument, leading to possibly dangerous clinical decisions. The basic principles of designing such surfaces are of course the same as those already invoked for finishing the surfaces of implants in biofluids. In one respect, the task of the bioanalytical surfaces of designer is easier, insofar as the contact duration is typically very short, maybe only a few seconds. The task is also easier insofar as many biofluids to be analysed may have been somewhat simplified, e.g. by centrifugation or filtration, prior to their introduction into the analytical device. Blood for example is almost invariably introduced as plasma or serum into such devices.

3.6. FURTHER READING

3.6

69

Further reading

A. Heberle (ed.), Human Protein Data, New York, Basel: VCH (1983). Ramsden, J.J. Bioinformatics: an Introduction. Dordrecht: Kluwer (2004). I. Roitt, Essential Immunology, 7th edn, Oxford: Blackwell, 1991.

70

CHAPTER 3. BIOFLUID COMPATIBILITY

entity

concentration

Proteins

function

µg/mL

albumin

general scavenger of smaller molecules

fibronectin

participates in thrombus formation

immunoglobulin G

part of the adaptive immune response

lysozyme

digests certain bacterial cell walls

transferrin

ion transport

Ions

mmol/L

Ca2+ Na+ K+ Cl− osmolytes

g/L

glucose

food

urea

metabolite

Cells

µL−1

erythrocytes

oxygen transport

neutrophils

destruction of microbial invaders

lymphocytes platelets

thrombus formation Table 3.3: The composition of human urine.

3.6. FURTHER READING

entity

concentration

Proteins (Mr )

71

function

mg/mL

albumin (66 500)

4

fibronectin

6

immunoglobulin G

general scavenger of smaller molecules

0.5

part of the adaptive immune response

(14 500)

1–3

digests certain bacterial cell walls

transferrin

0.1

iron transport

Ions

mM

Ca2+

0.5

Na+

120–170

+

30–35

lysozyme

K

Cl−

110–140

osmolytes

g/L

glucose

0.1

urea

0.2–0.3

food metabolite

Table 3.4: The composition of human tears.

72

CHAPTER 3. BIOFLUID COMPATIBILITY

Chapter 4 Cell and tissue interaction with surfaces Other than in the bloodstream, in which individual cells are circulating, cells in the body are present as a confluent mass (tissue), which may be soft (e.g. skin, tendon, pericardium, cornea) or hard (e.g. bone, dentine, cuticle). The cells constituting tissue are held together by the extracellular matrix, composed of proteins and glycoproteins. As living matter, tissue is more-or-less active (i.e. the metabolic rates vary enormously), modulating their environment according to the stimuli they receive. One of the most important of these modifications is the coating of an artificial implant with a layer of proteins, synthesized by the tissue cells and known as microexudate.1 Most medical implants are in intimate contact with human cells, and this contact needs to be considered very carefully (biocompatibility with hard and soft issues). An essential principle is that cells are almost never in direct contact with an artificial surface: there is almost always an intermediate layer of protein, and if the protein is not provided beforehand, the cells will typically synthesize and excrete it themselves before bonding to the surface. A second principle is that the behaviour of the cells is modified by their interaction with the surface, i.e. there is an active flow of information between the surface and the cell. Even devices in the bloodstream, albeit more significantly in contact with the blood, are typically partly embedded in the endothelial lining of the blood vessels. 1

G. Poste et al., The study of cellular “microexudates” by ellipsometry and their relationship to the cell coat, Experimental Cell Research, 78 (1973) 303–313.

73

74CHAPTER 4. CELL AND TISSUE INTERACTION WITH SURFACES

4.1

Principles of eucaryotic cell architecture

The eucaryotic cell is a small three-dimensional body, of irregular shape and typically about 10 µm in length, breadth and width. Its outer surface is comprised of a thin membrane, often called the bilayer lipid membrane, but in reality composed not only of lipids but also of proteins and glycoproteins embedded in the lipid bilayer (see Figure 4.1). Two very important functions of the membrane are: (1) to control the passage of ions, osmolytes and proteins in and out of the cell; and (2) to receive signals from the environment and to transmit them to the cytoplasm. CellMem Figure 4.1: Sketch of the so-called lipid bilayer membrane showing peripheral and integral membrane proteins and glycoproteins. H, lipid head-group; T, lipid tail; P, peripheral membrane protein; A, protein and ca, comprised of either an acyl chain, or hydrophobic amino acid residues, or both; E, extracellular domain of an integral membrane (glyco)protein; Tr, transmembrane portion of the integral membrane protein, typically comprising one or more of the alpha helices traversing the membrane; C, cytoplasmic domain, which may be simply a loop connecting the transmembrane portions.

The cytoplasm is a viscous, crowded aqueous medium. It is full of internal membranes, similar to the bilayer lipid membrane bounding the cell is a whole, and to which most of proteins are bound, some transiently and others more or less permanently. These internal membranes in turn enclose organelles, quasi-miniature cells such as the mitochondria (specialized in energy production and storage), the endoplasmic reticulum (specialized in protein “finishing”) and the nucleus, home to the DNA on which the genome is stored, and also containing the supramolecular “machinery” that transcribes the DNA into messenger RNA, which in turn is exported across the nuclear membrane into the cytoplasm and translated into protein by the ribosomes, large supramolecular complexes. The genomes of most eucaryotes are too large to be present as a single DNA molecule. Instead, the DNA is organized into several sections, each one containing many hundreds or thousands of genes, called chromosomes. Strictly speaking the chromosome, a highly condensed form of DNA, only exists prior to cell division, i.e. in populations of proliferating cells. Otherwise the DNA exists in a less condensed form suitable for templeting transcription.

4.1. PRINCIPLES OF EUCARYOTIC CELL ARCHITECTURE

75

DNA is complexed with proteins (histones) around which the DNA (150– 200 base pairs) is wrapped (these units are called nucleosomes). Accessibility to the enzymes involved in transcribing the genes into messenger RNA is controlled by conformational changes in these proteins, and by the DNAbased mechanisms of promotion described in the next paragraph. Promoters are groups of bases (nucleic acids) very close to the bases encoding the genes proper. Since transcription occurs in one direction, it is usually stated that the promoter sites are upstream of the gene. The transcription factor must bind, or be released from, the promoter site before transcription can begin. Sometimes a single site may control several successive genes. Transcription control can also be affected by the binding of small (typically fewer than 20 bases long) RNA molecules, called interfering RNA or iRNA. Obviously if a complementary strand of RNA binds to the DNA, transcription factor binding will be blocked. There also exists the possibility of chemically modifying certain bases, notably adenosine, by methylating them. This typically blocks expression. Methylation patterns are not inheritable in the conventional sense (i.e. by DNA replication), but are typically passed on to offspring when a cell divides, thus providing a mechanism for a kind of Lamarckian inheritance. The mechanism of separating the chromosomes prior to cell division involves some exceedingly complicated molecular machinery that is probably rather sensitive to damage. Although the cells of a healthy human being are diploid (i.e. they contain one set of genes from the father and one from the mother) and have a constant number of chromosomes (23), this pattern can be perturbed, most notably in cancer cells, in which there may be irregular numbers of many of the chromosomes, and many of the chromosomes themselves may be distorted (e.g. shorter than the standard length). This facilitates the rapid evolution so characteristic of cancers. It should be pointed out that many plants are polyploid: it appears to be a way of compensating for their limited abilities to evade environmental adversity, since they are not mobile. But animals do not appear to be as good at chromosome management, and polyploidy often has fatal consequences (in the form of cancer). This is important in the present context because tumorigenesis can be a response to a biomedical implant. The shape of the cell is controlled by the cytoskeleton, a network of proteins aggregated into fibres. The main structural fibre is the actin filament. Microtubules, made from the protein tubulin, are very active upon cell division, but also pervade the cytoplasm and other times and function as tracks on which special “locomotives” (kinesins) run, transporting molecules, especially proteins, from one part of the cell to the other.

76CHAPTER 4. CELL AND TISSUE INTERACTION WITH SURFACES What space is left between all these molecules is filled with water molecules, ions (mainly potassium, magnesium and calcium), small metabolites (e.g. adenosine triphosphate, glucose) etc. As the classic experiments by Kempner and Miller2 have shown, this medium has character of a hydrogel rather than ordinary liquid water. Figure 4.2 summarizes the picture. Cell Figure 4.2: Sketch of a eucaryotic cell.

4.2

Cell adhesion

Cells grown in a liquid culture medium will usually be spherical—as expected from a simple application of energy minimization and Laplace’s Law. This is of course a highly artificial state of affairs for most cells. A fascinating and complex sequence of reactions ensues when a cell “lands” on a solid substrate. The dramatic shape change occurs—from sphere to segment. The segment is of course a Platonic ideal: the cross-section is irregular, and the plan even more so. During this change the cell-substrate contact area increases from zero to tens of square micrometres. Formally the cross-section of a sphere, of radius r0 , is A(z) = πz(2r0 − z) ,

(4.1)

A(z) = πz(2r − z) .

(4.2)

and that of the segment is

The height h and cross-sectional area a are fixed by the relations a = πh(2r − h)

(4.3)

and the expression for the volume V : V = πh2 (r − h/3) 2

(4.4)

Kempner E S and Miller J H The molecular biology of Euglena gracilis IV. Cellular stratification by centrifuging 1968 Exp cell Res 51, 141–149; Kempner E S and Miller J H The molecular biology of Euglena gracilis V. Enzyme localization 1968 Exp cell Res 51, 150–156

4.2. CELL ADHESION

77

which may be taken to be the same as that of the initial volume of the spherical cell, (4.5) V0 = 4πr03 /3 . Spreading can take place in minimal medium under which no growth of the cell occurs. In general however the cell might well augment its volume. Initially, exploratory filopodia protrude from the cell, driven by localized actin polymerization. Sophisticated topographical, mechanical and chemical information is relayed from specialized receptors to the nucleus of the cell, and an appropriate response is generated: either to continue spreading, or to abandon the attempt. Science still has a very incomplete picture of this process. Figure 4.3 illustrates the spread cell. CellSpread Figure 4.3: Sketch of a spread eucaryotic cell (in cross-section). On the left the overall shape, on the right detail of the cell-substrate interface. Quite a lot of detail is known about the molecular constituents of the filopodia.

The focal adhesions at the extrema of the filopodia are typically terminated by large transmembrane proteins in the cell wall called integrins. The integrins can specifically bind to certain combinations of amino acids such as arginine-glycine-aspartic acid (RGD). The cytoplasmic side of the integrins is connected to the actin cytoskeleton by various proteins that may include MARCKS, talin and vinculin (their detailed roles are not understood at present). Some important circulating cells such as leucocytes have very specialized cell-substrate interaction mechanisms because their role is centred on their ability to recognize particular zones of certain tissues. These interaction mechanisms are based on complementary hydrogen bond-based biochemical attraction between molecules in the roving cell wall and in the target tissue cell wall. In the normal healthy state of the body leucocytes are travelling along with the blood flow and are not even able to come into contact with the walls of the blood vessels (venule). Only some perturbation (such as that caused by physical damage, or inflammation) of the venule reduces the hydrodynamic forces (principally the Magnus force) and allows them to reach the endothelium. There, short-range molecular recognition between large trans-membrane glycoproteins such as integrins embedded in the leucocyte cell surface and other large trans-membrane glycoproteins called selectins

78CHAPTER 4. CELL AND TISSUE INTERACTION WITH SURFACES embedded in the cells constituting the walls of the endothelium increases the attraction between the leucocyte and the endothelial surface. This adhesive force combined with the hydrodynamic force pushing the leucocyte along with the blood flow causes the leucocyte to roll along the venule surface, whereupon it can explore its detailed chemistry more readily. At the site of acute damage or inflation integrins are likely to be encountered in the endothelial cell membranes, and stronger adhesion between the leucocytes and endothelial integrins prevents the leucocytes from rolling further, whereupon they push themselves between adjacent endothelial cells and penetrate into the venule wall.3 Table 4.1 lists some of the molecules that are known to be expressed at the surfaces of the leucocyte and the inflamed endothelium during the process.4 The mutually interacting selectins and mucins are typically associated with transient cell-cell adhesion, whereas cadherins (which engage in homotypic binding) are associated with stable cell-cell adhesion. As Table 4.1 shows, the integrins are involved in both transient and stable adhesion. They typically interact with extracellular matrix proteins such as fibronectin. “Cell adhesion molecule” (CAM) is the umbrella term that covers most of these molecules. They are typically large glycosylated molecules, constructed using a modular principle. Hence for example the integrins are typically given labels such as α4β1 or α4β7. The notation is fairly unsystematic at present: some molecules have multiple names, for example LFA-1 is actually integrin α2β2. Single molecule techniques are now yielding detailed information on individual ligand-receptor interactions of the cell adhesion molecules,5 but such results should always be seen in the context of the fact that these molecules almost never act in isolation, but in clusters, the clustering itself being an adaptive response to an initial, possibly nonspecific, adhesion event.

4.3

The adhesion of cells to artificial surfaces

Cells have developed specialized molecular architecture primarily in order to engage in interactions with other cells, as described in the previous section for leucocyte homing. A first approach towards investigating the interactions of 3

Bruinsma, R. Physical aspects of adhesion of leucocytes. In: Physics of BioMaterials: Fluctuations, Self-Assembly and Evolution, T. Riste and D. Sherrington, eds, pp. 61–101. Dordrecht: Kluwer, 1996. 4 D. Dunon et al., To stick or not to stick: the new leucocyte homing paradigm, Current Opinion in Cell Biology 8 (1996) 714–723. 5 E.g. A. Pierres et al., Determination of a lifetime and forced dependence of interactions of single bonds between surface-attached CD2 and CD48 adhesion molecules, Proceedings of the National Academy of Sciences of the USA, 93 (1996) 15114–15118.

4.3. THE ADHESION OF CELLS TO ARTIFICIAL SURFACES Stage

leucocyte molecules endothelial molecules

Rolling

selectins

?

integrins

CAM

CD44

HA

?

chemokines

PECAM-1

PECAM-1

LFA-1

CAM-1

integrins

CAM

LFA-1

CAM-1

integrins

CAM

Triggering Tight adhesion Penetration

79

Table 4.1: The molecules expressed during the successive stages of leucocytes homing in from the bloodstream to the vascular endothelium. Where ligand abbreviations: CAM, cell adhesion molecule; HA, hyaluronic acid; LFA, leucocyte function associated molecule. cells with artificial surfaces would be to inquire to what extent features of the artificial surfaces correspond to features of the living surfaces. A discussed in the following section, bio-surfaces may comprise both the surfaces of the living cells, and the basement membranes constituted from extracellular matrix molecules on which cell assemblies are constructed. According to the principles elaborated in Chapters 1 and 2, hydrogen bond donors and acceptors are the main binding-determining features. A major question concerns the importance of the exact arrangement of these donors and acceptors (i.e. Lewis acids and bases) on the substrate. The surface of the cell, constituted by a cell membrane, in semifluid, and the arrangement of Lewis acids and bases can adapt to its environment. Therefore, one might reasonably postulate that the exact geometrical arrangement is not of importance. We have, however, already shown (Chapter 2, §2.1.2) that a mean-field approach inadequate for predicting protein adsorption. Therefore, it is evident that some statistical parameter, giving information intermediate between the global average of the mean field and the exact enumeration of the atomic coordinates of each constituent, of the molecular arrangement of acids and bases need to be found, but there is no clue as to what that parameter is.6 We are still at 6

The ability of “floppy” antifreeze glycoproteins to selectively absorb on particular ice crystal facets is a further piece of evidence in favour of the intermediate statistic.

80CHAPTER 4. CELL AND TISSUE INTERACTION WITH SURFACES the stage of laboriously accumulating data. Furthermore, as already emphasized, the ubiquity of water must play a decisive role, especially considering that the water structure itself is modified by the surfaces in its vicinity. The artificial materials from which biomedical device are made are either pure elements are compounds, or crystalline or amorphous. Especially in the case of thin films coating about device, the crystallinity or otherwise of the surface is not always known. Nevertheless, it can play a decisive role. In the case of implants made from pure elements, crystallinity may be of secondary importance. However, only gold and platinum, among the common implant materials—and even they are quite rarely used—remain oxide-free. In the case of compounds, the exposed plane of a crystal of compound MX can be either solely M, solely X, or both. M may well be a Lewis acid, and X a Lewis base: therefore it is vital to know what the plane is. Fundamental information is therefore to be gained from cell growth experiments on a single crystals. Intriguing results have been obtained by Hanein et al.7 They grew epithelial cells on different faces of calcium carbonate tetrahydrate, which mainly differ in the surface distribution of their lattice water molecules. Initially (within the first few tens of minutes) cells only attached to the (011) faces, where they started to spread, but after 72 hours they were all dead, yet well-spread living cells were present on the (101) faces. No deeper explanation can as yet be offered for this striking phenomenon.

4.4

Principles of tissue architecture

Tissue is defined as an aggregate of individual cells. It is typically classified into four basic types: Epithelial tissue: the epidermis of the skin, the mucosa of the gut, follicles of the glands, and the retina, for example; Connective tissue: the dermis of the skin (loose connective tissue), tendons (dense connective tissue), adipose tissue, supportive tissue (bone and cartilage); haemopoietic tissue (bone marrow) is also included here; Muscle: smooth (artery), skeletal (limbs, lungs), cardiac (heart); Nerve: brain, central nervous system, peripheral nervous system. There are many contrasts in tissue, such as dry (skin) or moist (the mucosae, the endothelia, i.e. inner linings of the circulatory system); soft 7

Selective interactions of cells with crystal surfaces, Journal of Cell Science, 104 (1993) 275–288.

4.5. ASSIMILATION

81

(e.g. cartilage) and hard (e.g. bone); etc. Adjacent cells in tissue are typically connected by channels in their respective cell membranes formed by proteins called connexins, allowing them to exchange a certain amount of material and, hence, information rapidly and fairly non-specifically. Many tissues are laid down on so-called basement membranes composed of a mass of extracellular matrix (ECM) proteins such as collagen, fibronectin, laminin and tenascin. Many of these ECM proteins have specific ligands for the integrins in the tissue cell surface, such as the RGD amino acid triplet (arginine-glycine-glutamic acid).

4.4.1

Inflamed tissue

Inflammation is an important general response of tissue to injury, which may be mechanical or cytotoxic, caused by chemicals or microbes. Most kinds of injury compromise the integrity of the tissue cell-cell contacts, hence blood plasma will automatically penetrate into the tissue. Leucocytes are also rapidly accumulated by the mechanism mentioned in the previous section. This phase is called acute inflammation. It is followed by chronic inflammation during which repair is attempted. Microbes and irreparably damaged tissue cells may be killed and decomposed, and residues (including undecomposable inorganic matter) transported away; any new tissue required will typically also involve the formation of new extracellular matrix (fibrosis) and new blood vessels (angiogenesis). If the injury is due to the presence of a foreign substance that can neither be decomposed nor removed, a persistent type of inflammation occurs.

4.5

Assimilation

All space within the body of an organism is filled; therefore the penetration of a biomedical device is almost inevitably going to result in injury. The response of the body to the intrusive device can take one of three general routes: 1. Perfect assimilation involves tricking the cells in contact with the device into thinking that the device is a bona fide part of the body. 2. Imperfect assimilation leads to fibrous encapsulation, typically avascular, of the intrusive device; this means that it is surrounded by protein such as collagen, in order to isolate it from the living matter.

82CHAPTER 4. CELL AND TISSUE INTERACTION WITH SURFACES 3. Non-assimilation means that the intrusive device remains as a persistent source of ongoing inflammation, that can usually only be brought to an end by removing the device. Although (2) might seem to be more desirable than (3), fibrous encapsulation often leads to problems after some time has elapsed. It is particularly undesirable in the case of prostheses with a mechanical function, such as a hip replacement, because the prosthesis is then less firmly bound within the body than it would be if it were in intimate contact with the body tissue, i.e. with its cells. It may move around with its capsule, causing mechanical insult to the cells surrounding it. As is well-known, all cells within the body (except the red blood cells) share the same DNA, but the more differentiated and specialized a cell has become, the less it is possible for it to express proteins uncharacteristic of its specialized phenotype, even though it carries the genes for those proteins, and the less it is possible for it to proliferate (i.e. divide). This is currently one of the greatest challenges to biological science: to be able to control cell differentiation. From the initial embryo at the birth of an organism onwards, the cells progress generally irreversibly along a path of increasing specialization. Stem cells may be totipotent, i.e. able to become any other kind of cell, or pluripotent, i.e. with the ability to become one of several specialized types. The more differentiated a cell is, the less possible it is for it to proliferate. That is why the repair of specialized tissue generally needs an influx of cells higher up in the potency hierarchy. The progressive specialization of cells is a typical epigenetic phenomenon. The cells are “thinking”, i.e. processing information received from their environment, living and nonliving. The internal clocks are also presumably providing information, in ways about which very little is known at present. The ability of cells to become assimilated with an implant depends very strongly on the information received and how it is processed, which is itself dependant on the informational environment. The investigation of assimilation is obviously crucial to the successful implementation of biomedical devices. Nevertheless, it is only relatively recently that a start has been made on establishing systematic correlations between biomedical surface attributes and cell physiological response. The main difficulty is perhaps the richness of “cause” (all conceivable morphological, mechanical and chemical attributes of the surface) and “effect” (all conceivable responses at proteomic, metabolomic and cell-morphological levels, all of which are interrelated). Add to this the fact the the effects are changing the nature of the cause, and one can understand why so little progress has been made. The sort of investigation that one might conceive is to allow a

4.6. BACTERIAL INFECTION OF IMPLANTS

83

certain cell type to grow on a series of surfaces of different mean roughnesses, or different chemical compositions, and to observe differences in protein expression and rate of morphological change.

4.6

Bacterial infection of implants

Bacteria are endemic within our world, indeed the human being is host to more bacterial cells (mainly in the intestine) than its own cells. The body is however well-tuned to finding and eliminating any foreign bacteria that penetrate inside it. The air and the food that we eat and drink is full of bacteria, many of which are trapped and neutralized before they even penetrate within the body, but for any that do succeed in entering the bloodstream, the chance of survival (i.e. before they are destroyed by the immune system) for a given duration is a random variable, and while they are still freely circulating, should they encounter a foreign surface, such as that of a biomedical device, they may well find it a congenial place at which to linger. Most of the bacteria able to survive in the bloodstream are able to rapidly proliferate. If one does adhere to a biomedical surface, and begins to proliferate (surrounded by an abundant source of nutrients at a congenial temperature), it will rapidly build up a little colony that will suitably protect itself from aggressive immune cells by erecting a composite structure in which the cells are embedded in a robust matrix constructed from extracellular proteins, glycoproteins and polysaccharides synthesized by the bacteria, forming a biofilm (see below).

4.6.1

Bacterial structure and dynamics

Bacteria, like eucaryotic cells, are also envelopes of cytoplasm surrounded by a membrane. Most bacteria are substantially smaller than eucaryotic cells, typically 200 nm long (although there is great variety). Different types (the word “species” is not generally used to describe different kinds of bacteria, because of the promiscuity with which bacteria can exchange DNA with one another) are characterized by different shapes: most typically spheres and spherocylinders (rods). The membrane has the same basic structure of a lipid bilayer with embedded proteins, but in the case of most bacteria is encapsulated within a tough polysaccharide wall that may be tens of nanometres thick. Within the membrane, the bacterial cytoplasm is also very crowded, but there are no internal membranes. Furthermore, there is no cytoskeleton. One of the most important features that distinguishes bacteria from eucaryotes is the genome plasticity. The gene management of bacteria is anyway

84CHAPTER 4. CELL AND TISSUE INTERACTION WITH SURFACES much more straightforward than that of eucaryotes, since mostly they do not engage in sexual reproduction; the DNA (present as a single strand) is simply replicated and the cell divided into two. Bacteria are quite susceptible to viral infections, involving the import of a significant quantity of foreign DNA, which may be incorporated into the bacteria’s own DNA, and similarly other bacteria can provide DNA. This means that bacteria can evolve for more rapidly than might be expected from the occurrence of random mutations in their own DNA, arising through the usual mechanisms of mis-copying or chemical damage.

4.6.2

Biofilms

Biofilms are frequently formed at solid/liquid interfaces by initially planktonic (freely swimming) bacteria settling on the solid surface. Biofilm formation is a very effective survival strategy: it is well known that bacteria are far more resistant to many different kinds of attacks, including powerful antibiotics, when living in a biofilm compared with the planktonic state. Part of this resistance is due to the trivial reason of retarded mass transport within the biofilm. It has also been suggested that bacteria living in a biofilm grow more slowly than their planktonic counterparts, and hence their uptake of antibiotics would be lower. But the most interesting and important difference is that the patterns of protein expression by the bacteria in the biofilm are very significantly different from those typical of planktonic bacteria.8 If a biofilm is populated by several types of bacteria, even without the possibility of DNA exchange between them (which is also very likely to occur), we have a kind of superorganism that is more capable of resisting external attack than the organisms individually.

4.6.3

Bacterial adhesion

Since bacteria generally have a rather rigid shape, there has been quite a lot of attempts to predict bacterial adhesion using the surface tension notions described in Chapter 1. At the time of writing however, this field is in a rather unsatisfactory state, both from the experimental and theoretical viewpoints. On the experimental side, it is still difficult to obtain reliable values for the interfacial energy components of a bacterium. Even when this data is available, there are usually insufficient experimental details given in the reports to allow the data to be applied with confidence for calculating 8

Vilain, S. et al., Bacterial proteome: homogeneity of versatility? J. Proteome Res. 3 (2004) 132–136.

4.7. FURTHER READING

85

the free energy of interaction of a bacteria with a surface, because bacteria are changing their phenotypes throughout their lives, and the precise state of relevance to the problem at hand is usually unknown.

4.7

Further reading

B. Alberts et al., Molecular Biology of the Cell, New York: Garland, 1994. R. Bongrand et al., Physics of cell adhesion, Progress in Surface Science, 12 (1982) 217–286. K. Burridge et al., Focal adhesions, A. Rev. Cell Biol. 4 (1988) 487–525. J.G. Edwards, The biochemistry of cell adhesion, Progress in Surface Science, 13 (1983) 125–196. R.J. Ellis, Macromolecular crowding: obvious but underappreciated, Trends in Biochemical Sciences, 26 (2001) 597–604. M.A. Hubbe, Adhesion and detachment of biological cells in vitro, Progress in Surface Science, 11 (1981) 65–138. G.H. Pollack, Cells, Gels and the Engines of Life. Seattle: Ebner (2001). S.A. Safran et al., Physics of cell elasticity, shape and adhesion, Physica A, 352 (2005) 171–201.

86CHAPTER 4. CELL AND TISSUE INTERACTION WITH SURFACES

Chapter 5 Lubrication and wear Friction is no less important in biology than it is in mechanical engineering. Whenever two surfaces are in relative motion to each other, friction is likely to occur. This not only causes heating, which may facilitate chemical transformations, but also mechanical damage, most notably wear, in which material is detached from the surfaces being rubbed together. This is obviously likely to weaken the objects whose surfaces are being rubbed, and in most biomedical devices the detached material, in the form of small particles, becomes dispersed within the body, where it can have very undesirable biological effects The negative effects of friction may be alleviated by lubrication. Most mutually rubbing surfaces in biology—the ball and socket joints, for example— are lubricated by special liquids such as synovial fluid. Artificial replacement joints may be designed to be lubricated by material infiltrating from the natural environment, or may be made from ultralow friction materials not requiring lubrication.

5.1

Fundamentals of tribology

Tribology is the study of rubbing, i.e. friction, and has also come to include the sciences of lubrication unaware of rubbing surfaces with its field. Not surprisingly, in view of its tremendous importance in machinery, friction was investigated quite systematically at an early stage of the Industrial Revolution; important basic laws were formulated starting more than 200 years ago, such as those of Coulomb (1785) and Morin (1833). 87

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CHAPTER 5. LUBRICATION AND WEAR

5.1.1

Dry friction

The fundamental law of friction defines a friction coefficient µ (and the corresponding angle φ from the minimum force S that is required to shift an object of weight W (also equal to the reactive force of the surface acting on the object) µ = S/W = tan φ (5.1) (see also Figure 5.1). When the object is initially at rest, this is known as the static friction coefficient µ0 . It was an early and important result of Amonton that µ did not depend on the contact area of objects, but only on its weight (or on the applied load). It was also known (Euler) that the friction diminished the motion, i.e. µ0 > µ. The beguiling simplicity of the early laws formulated by Coulomb and Morin were rightly criticized by Hirn, who asserted that they were much too simple to be exact.1 Furthermore, they were not applicable to the much more important practical cases of friction mediated by lubricants (see below). fric1 Figure 5.1: Dry friction.

Most of the mystery of friction was resolved when it was realized that the detailed morphology of the rubbing surfaces at all length scales needs to be taken into account. Bowden and Tabor formulated their law of the friction force as FR = τ AR (5.2) where τ is the shear strength and AR is the real area of contact between the two surfaces. Typically this real area is several orders of magnitude less than the nominal area. This law finally put paid to Coulomb’s notion that there was no friction between perfectly smooth surfaces. Furthermore, since the asperity could be modified by wear taking place during the rubbing, it is really necessary to define friction force as a functional of the lateral rate of displacement, x(t). ˙ It is also accepted that most of the work done against friction is liberated as heat; only a small fraction is retained as strain energy in the lattice. 1

See for example G.A. Hirn, Note sur les lois de frottement, C.R. 99 (1884) 953. But this memoir was published about 40 years after Hirn’s experiments were initiated, due to some kind of dispute between him and members of the Acad´emie des Sciences, the details of which have been lost.

5.1. FUNDAMENTALS OF TRIBOLOGY

89

Knowledge of friction has continued to improve in recent years due to the advent of new experimental techniques, most recently and most notably scanning probe microscopies that can be used to investigate friction down to nanometre length scales. These measurements have also shed light on the stick-slip phenomenon, due to the elastic deformation of individual components of the asperity.

5.1.2

Lubrication

The enormous dissipation of heat due to friction in early machinery provided a powerful impetus to the development of lubricants. Lubrication refers to the modification of friction due to the presence of additional matter between the rubbing surfaces. Initially it was invariably achieved by adding liquid lubricants such as oils of various kinds. The dependence of friction on load becomes more complex, as sketched out in Figure 5.2. The shape of the curve, for two surfaces moving with a relative velocity v, is given by the equation F = FC + (FM − FC ) exp(−v/vS ) + µv v (5.3) where FC is Coulomb’s (1785) coefficient, FM is Morin’s (1833) static or breakaway friction, vS is Stribeck’s (1902) empirical characteristic velocity, at which the opposing asperities disengage, and the final term is due to Reynolds (1866) and accounts for the viscosity of the lubricant (µv is the viscous friction coefficient). fric2 Figure 5.2: Lubricated friction. At very low velocities one refers to boundary lubrication, or nanofriction, because the lubricant film is less than 100 nm thick. More exactly, the ratio h/R of the thickness h of the lubricating film to the roughnesses R of the surfaces is low. Around the minimum of the curve one has mixed film (elastohydrodynamic) lubrication. At very high velocities hydrodynamic lubrication dominates: the friction is determined merely by the viscosity of the lubricant; the surfaces are no longer in contact with one another. More recently, efforts have been made to develop fixed coatings of highly deformable organic molecules such as neutral or charged polymers grafted to the rubbing surfaces as a substitute for adding liquid lubricants.2 This 2

U. Raviv et al., Lubrication by charged polymers, Nature (Lond.) 425 (2003) 163.

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appears to be one of nature’s strategies. In the context of biomedical surfaces, one can hardly avoid lubrication since (aqueous) fluids are ubiquitous. Many of these fluids have extremely complex compositions (Table 5.1 gives the composition of the major components of synovial fluid), and it is by no means obvious that they all act to reduce friction; on the contrary, certain components absorbing onto the rubbing surfaces may actually increase the friction, in a desirable and controlled way (cf. the brakes on a motorcar!). In the living body, it appears that the composition of lubricants is carefully controlled. One of the most important (and ubiquitous) components is the large glycoprotein mucin. Even coating mutually rubbing surfaces by simple physisorption of mucin gives dramatic reductions in sliding friction.

entity

concentration

Proteins

µg/mL

albumin

4–10

fibronectin

0.15

immunoglobulin G

1

function general scavenger of smaller molecules part of the adaptive immune response

lysozyme

absent

digests certain bacterial cell walls

transferrin

absent

iron transport

Ions

mM

Ca2+ 2 Na+

135

K+

4

Cl−

100

osmolytes

g/L

glucose

absent

food

urea

absent

metabolite

Table 5.1: The composition of synovial fluid (major (by quantity) components).

5.2. FRICTION IN HIP ENDOPROSTHESES

5.2

91

Friction in hip endoprostheses

Figure 5.3 shows actual measurements of the friction coefficients of typical pairs of materials used for the combination of head and acetabular cup. One notes that the friction coefficients are not constant in time; the results of wear are indeed often complex and unpredictable. 17MA-2 Figure 5.3: Friction coefficients of (from top to bottom and five main): α-alumina ceramic/polyethylene; Co-Mo-Cr alloy/polyethylene; and monocrystal sapphire/polyethylene.

5.3

Possible physiological effects of particles released as a result of wear

Whereas the materials (metals, ceramics and polymers) used for those biomedical implants that are subject to wear (i.e. load-bearing joint replacements) or deployed because of their desirable mechanical properties (mainly strength) and resistance to corrosion, a quite different set of properties are relevant when considering the fate of particles released by wear into the body. Susceptibility to corrosion actually becomes an advantage (unless the material is toxic), because the particles will simply then dissolve away. Mechanical strength is irrelevant. One is mostly concerned with interfacial reactions taking place at the particle surface. In reality, mechanical abrasion and chemical corrosion combine to erode surfaces of initially pristine implants. Figure 5.4 shows an example of the results of 10 years of use of a hip endoprosthesis. 17MA-1 Figure 5.4: Wear of the metallic head and ferulene bush of a normal endoprostheses after 10 years of operation.

The most obvious consequence of wear is that the joint components diminish in size and become loose. Physiologically however the most serious consequences arise from the material that is detached and released into the rest of the body. Although synthetic polymer prostheses wear rather badly,

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at least the particles they release tend to be a rather large, migrate a limited distance from the joint, and although they may cause some inflammation, the physiological response tends to be limited and localized. Potentially the most dangerous materials detached from bulk implants are the metal oxides. In practice this not only means debris from ceramics, but also from primary metal implants (Figure 5.5), because they are almost invariably coated with native oxide. Many of these ceramics wear much less than the synthetic polymers, but concomitantly the particles they release tend to be much smaller— nanoparticles. They migrate with ease throughout the body, (Figure 5.6) even into the brain. Metal oxide monocrystals, e.g. of alumina (sapphire) or zirconia, wear even less, but the released particles may be even smaller and more dangerous. There is still a considerable research deficit regarding the biological consequences of nanoparticles within the body. Revell-1a Figure 5.5: Electron micrograph by SEM of Co-Cr particles (0.3–5 µm) retrieved from tissues adjacent to a human implants at revision surgery.

Revell-1b Figure 5.6: Electron micrograph by ESEM of a microparticle (centre) and nanoparticles (top left, bottom centre) in the liver of a rabbit following implantation in the femoral bone marrow. The particles used are those shown in the previous figure.

These metal oxides are typically wide band gap semiconductors (Figure 5.7). Electrons in the conduction band are typically strong reducing agents, and defect electrons (positive holes) in the valence band strong oxidizing agents. Electrons and positive holes can be created by chemical, thermal or optical excitation. sc Figure 5.7: Sketch of the energy levels of a semiconductor. CB, conduction band; VB, valence band; ED, electron donor level (due to an impurity or defect).

5.3. POSSIBLE PHYSIOLOGICAL EFFECTS OF PARTICLES RELEASED AS A RESULT OF WE

5.3.1

Physical properties of (nano)particles

The most important relevant physical attributes of particles likely to be released from rubbing surfaces are (i) the very large surface area for a given mass; and (ii) their high curvature.

5.3.2

Chemical properties of (nano)particles

If particles are able to accumulate electrons in their conduction band, or positive holes (defect electrons) in the valence band, or both, biologically dangerous reactions may occur. For titania in a humid environment, a typical set of reactions would be + O2 −→ O•− 2 •− + O2 + H −→ HOO• ⊕ + OH− −→ • OH

(5.4) (5.5) (5.6)

• The radical species O•− 2 and OH are very powerful reducing and oxidizing species, respectively. Such reactions have recently been given prominence in connexion with photocatalysis (photoelectrochemistry), in which the electrons and positive holes are generated by the ultraviolet portion of sunlight. Particles that happened to migrated to the skin could be subject to irradiation and hence radical generation via optical excitation. In other cases the chemical and thermal routes to radical generation are omnipresent. Because of their large surface area, nano particles can be considered essentially of macromolecules and can actively participate in the chemistry of the environment. This includes exchange of ions with their lattice is well as the selective adsorption of material on their surface. Most oxides are known to be good protein adsorbents from studies using bulk material, or thin films. Nanoparticles migrating around the human body are likely to be accompanied by a retinue of many different proteins and other macromolecules, transporting them into places where the body would not normally expect to find them. The biological effects of this promiscuous protein transport are, as yet, little known and even less understood. In compound materials, formed from a cation C (metal) and an anion A (oxygen), the occurrence of defects is governed by the constraint of the need to maintain electrical neutrality, i.e. vacancies must be compensated by interstitial ions of the same sign (Frenkel type) or by vacancies of the

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opposite sign (Schottky type).The following relations are possible C◦ + C0 null A0◦ + A null C◦ + A0◦ CA

(5.7) (5.8) (5.9)

using Schottky’s notation—C◦ is an interstitial cation, and C0 is a cation vacancy, ‘null’ signifies the undisturbed lattice, and CA represents the addition (at the surface of the nanoparticle) of a new lattice molecule. A full circle is used to represent the substitution of a native ion with a foreign one, e.g. X• (C) would be a divalent cation X on a cation lattice site, Cu× • (Ag) would be (mono)valent copper on a silver lattice site in (say) AgCl, etc. The superscript respresents the electrical charge of the defect, relative to the undisturbed, overall electrically neutral lattice: 0, , × represent negative, positive and neutral (zero) excess electrical charge. The above three equations can be used to derive some more: null C0 + A + CA C◦ CA + A A0◦ CA + C0 .

(5.10) (5.11) (5.12)

To each of these there corresponds a mass action law, i.e. for the last three, writing x for the mole fraction, identified by its subscript: xC0 xA = K xC◦ /xA = K1 xA0◦ /xC0 = K2

(5.13) (5.14) (5.15)

where the K are the equilibrium constants. The defect concentrations are supposed to be sufficiently small so that the chemical potentials µ depend on the logarithm of x, e.g. for vacancies: µ − E = −RT ln x

(5.16)

where E is the energy to annihilate one mole of vacancies. Of course the usual electroneutrality condition holds: xC◦ − xC0 = xA0◦ − xA .

(5.17)

This formalism can be powerfully applied to the effect of impurities on defects. Examples: (5.18) NaCl Na× • (Ag) + AgCl

5.4. FURTHER READING

95

(silver chloride doped with sodium chloride): here there would be practically no change in defect concentration; but if the doping is with divalent cadmium, there would be a significant increase in silver vacancies and a decrease in interstitial silver ions: CdCl2 Cd• (Ag) + Ag0 + 2AgCl

(5.19)

Ag◦ + CdCl2 Cd• (Ag) + 2AgCl .

(5.20)

and In these mixed phases, the electroneutrality condition becomes xC◦ − xC0 = xA0◦ − xA + y

(5.21)

where y is the lattice concentration of the dopant. This relation, together with the three eqns 5.13, can be used to derive explicit relations for the concentration of each of the four defect types on y. In very small particles, a significant fraction of the atoms of the particle may be actually surface atoms. Because of their different bonding, the surface atoms may be sources or sinks of defects, thereby perturbing the equilibria (5.7). The adsorption of impurities at the surface of a nanoparticle will generate defects. For example, adsorption of two monovalent Ag+ onto the surface of a ZnS nanoparticle will require the formation of a zinc vacancy in the cation sublattice. The environment of the nanoparticle (the matrix) thus plays a crucial role in determining its properties. The biofluid environment of nanoparticles is abundant in certain ions, especially calcium (divalent), sodium and potassium (monovalent). More elaborate nanoparticles made from multiple elements may arise from metal alloys (details will depend on the phase diagrams of the alloy and its oxides), which may themselves be pure elements or compounds. The particles themselves may either be homogeneous, or segregated into zones. Finally, the small particles may simply dissolve and engender the toxic effects associated with their metal ions (cf. §2.4.2).

5.4

Further reading

F.P. Bowden and D. Tabor, Friction and Lubrication of Solids. London: Oxford University Press (1954). B.N.J. Persson, Sliding Friction: Physical Principles and Applications. Revell, P.A. (2006). The biological effects of nanoparticles. Nanotechnology Perceptions, 2, 283–298.

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CHAPTER 5. LUBRICATION AND WEAR

Chapter 6 Surface generation, engineering and coating This chapter deals with the modification of the texture (morphology and chemistry) of the surfaces of biomedical components and devices by machining, engineering and by coating. The selection of materials for biomedical implants in still primarily made on the basis of bulk properties and cost. Traditionally, ideal materials were considered to be totally inert with respect to the living environment, both their bulk and a fortiori their surfaces. We now know that this is a utopian and unrealistic view: the intrusion of any foreign substance into the body, no matter how smooth and chemically unreactive, will evoke some kind of biological reaction. The very fact that it occupies space will provoke some internal rearrangement in its immediate vicinity. Table 3 shows that titanium is overwhelmingly the most popular material used in biomedical implants. Its mechanical properties are very good, and its cost is acceptable. Remarkably enough, titanium dioxide, with which the bulk metal is almost invariably coated, is an extremely biocompatible material. The origins of this biocompatibility lie in its surface energetics (see Table 2.1.1). Furthermore, it so happens that titanium and its oxide are extremely corrosion-resistant with respect to the environment encountered with in the human body. Another example of a material that appears to be highly biocompatible is sapphire (single crystal alumina). In such cases, machining is sufficient to finish their surfaces; some post-shaping process might only be required in order to deliberately roughen the surface with an elaborate multiscaled texture, for example. In other cases, the material selected may be quite bio-incompatible, and therefore the surfaces in contact with the body must be modified in some way. The term “surface engineering” refers to the post-shaping mechanical, physical or chemical modification of the surface by adding or removing se97

98CHAPTER 6. SURFACE GENERATION, ENGINEERING AND COATING lected atoms, for example. The term “coating” refers to the addition of a thin film of a different material to the surface. Ultimately, the entire manufacturing process is subject to a principle of “least effort”: what is the fastest and cheapest route to attaining the desired attributes? In most cases, however, the problem cannot be formulated in terms of a cost function that simply has to be minimized. There are multiple competing or conflicting objectives; multi-objective optimization implies finding any one of some or many possible solutions that lie on the so-called Pareto front. Some compromise is inevitable: the task is to find the best possible compromise. On the Pareto front, one cannot improve any variable without causing a deterioration of another variable. Given the multiplicity of solutions on the Pareto front, the final choice is essentially a matter of aesthetics (including tradition and the personal preference of whoever is in charge of the particular process being optimized). In the case of most biomedical devices, not only is the manufacturing route complex, but the finished product has to have many possibly conflicting attributes. Compromise therefore is inescapable. Whereas for traditional engineering objects, whether large (e.g. a bridge) or small (e.g. a microprocessor), a materials selection diagram is often an adequate design aid, the sheer complexity of the biomedical environment makes that approach inadequate. Indeed in the field it is not customary to use materials selection diagrams; on the other hand there is no other comparable objective procedure in customary use at present. A deeper discussion of design principles is beyond the scope of this book; nevertheless, we can point the way to the large body of work already carried out in the field of multi-objective optimization. Because of the sometimes huge multiplicity of parameters, stochastic approaches based on evolutionary computation (e.g. genetic algorithms) are often necessary. A particularly useful and important development in the field has been the emergence of interactive genetic algorithms, which enable qualitative knowledge and personal preferences rooted in intuition and experience to be introduced into the evolutionary process. In traditional fields of engineering, the surface very often played an aesthetic role. One thinks for example of the intricate texture of Japanese swords; in fact to the eyes of the connoisseur it gives an indication of the quality of the manufacture, and hence effectiveness as a weapon.1 Except in some rather minor fields, such as anti-reflexion glasses, the actual surface finish of a product hats typically received rather cursory attention. 1

See for example C.S. Smith, A Search for Structure, pp. 103 ff. Cambridge, Mass.: MIT Press (1981).

6.1. MACHINING

6.1

99

Machining

The basic processes used to impart form and structure to an object are cutting, moulding and casting, forming and joining. Each of them has particular implications for the state of the surface of the finished object. More modern methods include, specifically, grinding, the addition and removal of material using electricity in one form or another, and the group of processes known collectively as semiconductor processing. Assembly is a process in its own right, and as medical devices become more and more miniature, traditional manual assembly methods may no longer be feasible. At the nanoscale, selfassembly has the potential for becoming a largely autonomous process. In the case of composites, assembly typically constitutes a major part of the operations required to create a finished piece, but since composite are as yet used relatively little for biomedical devices, we shall not go further into the details of composite assembly. One of the most traditional operations is to machine shapes by removing matter from unformed blocks (cubes) of material. The essential processes of cutting and grinding have remained conceptually unchanged for many decades, even though they are carried out at much higher precision than formerly.2 In the case of devices that only remain in contact with the body for a very short time, it may be possible to impart the necessary attributes of the device entirely by the machining process. A good example of this is the manufacture of hypodermic needles. The required attributes are relatively simple: painless insertion and inertness with respect to the contents being injected. Careful design of the shape of the needle has succeeded in minimizing the painlessness of the final object. Nevertheless, to ensure that the cutting edge that makes the initial incision in the skin is as sharp as possible, electropolishing (see below) has been found to be more cost-effective than mechanical removal alone. One aspect that needs to be carefully considered is that any mechanical process involving the removal of material may leave residual surface stresses in the workpiece. They may be removed by e.g. a thermal annealing process. The effects of such a surface stresses on biocompatibility, for example the response of living cells, are as yet largely uninvestigated.

2

N. Taniguchi, Current status in, future trends of, ultraprecision machining and ultrafine materials processing them. Annals CIRP 32 (1983) 573–582.

100CHAPTER 6. SURFACE GENERATION, ENGINEERING AND COATING

6.1.1

Mechanical operations

Machining processes such as precision and now ultraprecision grinding, diamond turning and micromilling are used not only to create devices directly, but also for creating moulds and templates with which inexpensive replicas, typically from synthetic polymers, can be made. The number of biomedical applications that use ultraprecision micromachining technology is increasing rapidly for several reasons. Firstly, there is a requirement for miniaturization of surgical equipment and devices in order to achieve less invasive surgery. An example of a tissue removal tool designed to be operated through an endoscope is shown in Figure 6.1. The ability to achieve smooth surfaces and well-defined edges is critical for this application. Kern Figure 6.1: Tissue removal tool for use with an endoscope directly manufactured from a solid block in under 15 minutes by 5-axis machining (courtesy of Kern GmbH).

Secondly, there is a growing demand for microsized prosthetic devices, whose realization is only possible thanks to fairly recent advances in manufacturing technologies and materials for their production. Some of the smallest prostheses are produced by micro injection moulding, which requires precision and ultraprecision machining for the production of the corresponding moulds (Figure 6.2). Examples of implants are illustrated in Figure 6.3. Gottschlich Figure 6.2: Brass mould insert for the injection moulding of microfluidic devices produced by micromilling (Gottschlich).

cochlea Figure 6.3: Micro cochlear implants fabricated by injection moulding, compared with the head of a safety match (courtesy of Battenfeld). Thirdly, there is a growing realization that in order to achieve acceptable assimilation within the host, prostheses and their surfaces need to be individually tailored, including being endowed with appropriate nanoscale surface

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texture, to optimize properties and performance such as adhesion, wear resistance and interface stability, and to minimize the trauma of insertion. Surface characteristics are also important for the expanding technology of microfluidic devices, used for biological or chemical analysis, including point-of-care analytical devices. Specific applications for such devices and systems include blood analysis, cell sorting and characterization, glucose monitoring, genetic analysis and general lab-on-a-chip functions. The low cost dictated by the preference for using disposable devices in such applications makes moulding the preferred manufacturing route for microfluidic components. For achieving suitable characteristics and performance from a moulded microfluidic device, the level of precision with which surfaces and features are generated on the mould insert by a micromachining process is of great importance (cf. Figure 6.2). In micromachining, ions are also used for controlled erosion of material at the nanoscale. These processes, such as deep reactive ion etching (DRIE) and ion milling are familiar from semiconductor processing technology.

6.1.2

Laser micromachining

The laser is well adapted as the cutting tool for the production of microparts. The industrial manufacture of stents employs laser micromachining, typically nanosecond diode-pumped solid-state Nd-YAG lasers for NiTiNOL. Stainless steel stents are also laser-cut. Apart from shaping, the laser is also suited to the fabrication of surfaces with special rough features. In such applications, interference between the photons in intense pulsed beams causes localized melting in periodic patterns. The detailed mechanisms are not yet well understood. Figure 6.4 shows an example of such laser interference surface texturing in steel, and replicated in a synthetic polymer. The surface was irradiated with a pulsed femtosecond Ti:sapphire laser. Intriguingly, a two-dimensional pattern of cone structures with feature sizes ∼ 10µm emerged, on which there was a submicrometre superimposed structure with a ripple spacing was typically between 600 and 700 nm, just shorter than the wavelength of the femtosecond pulsed laser (800 nm). The ripple pattern was in fact found to be a general phenomenon occurring on many kinds of materials after irradiation with a few pulses.3 Further passes result in degradation of a ripple pattern, with the appearance of trenches perpendicular to the ripples. By applying even more pulses, highly chaotic structures emerge. 3

Groenendijk, M. and Meijer, J., 2006, Microstructuring using femtosecond pulsed laser ablation, J. Laser Applications, 18, 227235.

102CHAPTER 6. SURFACE GENERATION, ENGINEERING AND COATING GM Figure 6.4: (a) Laser interference surface texturing on a steel mould surface; (b) the texturing replicated in polypropylene (courtesy of Groenendijk and Meijer).

Some of the replicated structures turned out to be ultrahydrophobic. Intensive investigation of naturally ultrahydrophobic materials such as the leaves of various plants (e.g. the lotus, the nasturtium, the lupin, see Figure ??) have revealed that their surfaces are rough on the microscale with a specific microstructure, which in turn is covered with tiny wax crystals. The composite microstructured surface enhances the hydrophobicity of the wax. In consequence, such leaves contaminated by dirt are easily cleaned by rainwater: the water droplets are unable to wet the surface, and any dirt on the leaf sticks to the droplets and is carried away as they roll off the leaf. lupin Figure 6.5: Lupin leaves, showing ultrahydrophobicity.

It is a great advance to be able to mimic nature such that by design of appropriate micro and nanostructures, engineers are able to control wetting. Applications for such surfaces are manifold: apart from self-cleaning properties,4 which might be useful for working surfaces, floors and walls in hospitals. Other technical applications include control of (microfluidic) flows bioanalytical devices, friction control in lubrication, control of drop formation in emulsification, and wetting control of scaffolds for tissue engineering. Laser sintering is an additive processes used to consolidate ceramic powders. Such materials have potential use for bone replacement material. Originally developed as a rapid prototyping technology, it should be able to produce artificial bones customized relatively cheaply to individual patient requirements. Photochemical machining (PCM) has been used to etch complex structures into polymer sheets for the fabrication of minute prostheses used in the eye (Figure 6.6). 4

The ‘self-cleaning’ requires the addition of bulk water, provided by the rain in the case of outdoor plants, in order to remove the dirt. Other ‘self-cleaning’ materials incorporate photocatalytic nanoparticles (cf. §5.3.2) and generate dirt-destroying radicals upon illumination; in the most favourable case organic dirt is mineralized to CO2 .

6.1. MACHINING

103 Kapton

Figure 6.6: Lens haptics (holders for replacement eye lenses) etched from 0.125 mm thick Kapton (polyimide) film (courtesy of TechEtch Inc., Mass.).

Standard semiconductor processing techniques such as photolithography and chemical etching are used to produce structures in silicon such as “painless” injection needles (Figure 6.7) by anisotropic etching techniques that follow atomically-flat, slow-etching crystal planes. needles Figure 6.7: So-called “painless” injection needles machined in silicon using photolithography and chemical etching (courtesy of Georgia Institute of Technology, USA).

6.1.3

Electrochemical methods

Electrochemical machining (ECM) involves the passage of Faradaic currents. In the case of electroplating, metal is added to a substrate; the reverse process is called electropolishing, in which material is removed. It is very useful for reducing the surface roughness of needles and stents and for the production of specific texture (roughness) that may aid the assimilation of bone prostheses. Electrolytic photoetching is a hybrid electrochemical and photonic process employing ECM and photochemical machining techniques familiar from the semiconductor industry to dissolve metal through apertures defined by photoresist stencils. This is another useful technique for the fabrication of stents from metal foils. Figure 6.8 shows the surface finish on the machined NiTi sidewall. It should be noted that the thick titanium oxide layer on the foil surface that results from this process gives the material its corrosion resistance and prevents leaching of potentially toxic nickel into the body. EPE Figure 6.8: Polished sidewall of a NiTi foil produced by electrolytic photoetching.

104CHAPTER 6. SURFACE GENERATION, ENGINEERING AND COATING

6.2

Surface engineering

The definition of the processes dealt with in this section is those that are separately applied to an object subsequent to its basic fabrication, but excluding coatings, which are dealt with below in §6.3 As already emphasized, the interactions that occur at the surfaces of biomedical materials are critical to their performance, and it is therefore essential that surfaces are tailored to achieve optimum structure and composition in order to control both adsorption (especially of biomacromolecules, and of living cells) and degradation (such as leaching and corrosion) processes. A range of surface treatment and coating processes is available for enhancing the performance of biomedical materials. Surface engineering can be considered in terms of surface treatments (modification of a surface by changing its composition or microstructure or both) or surface coating (the addition of a distinct layer of material to the original surface). The surface engineered system can therefore be considered as a composite system, consisting of a near-surface modified region, which is optimized for the working environment, and the substrate or bulk material, which is usually selected based on mechanical and physical characteristics and manufacturability. The objective of creating such a composite system is that it will exhibit a performance superior to the coating or substrate alone. For biomedical applications such as implants and prostheses, the majority of coatings and surface treatments are applied to enhance corrosion or wear resistance, or to help with hard tissue compatibility by promoting bone growth.

6.2.1

Plasma modification

Plasma modification (e.g. through a dielectric barrier discharge in an inert gas) of polymeric materials such as polyamide (PA-6), poly(vinyl chloride) (PVC), poly(ethylene terephthalate) (PET), PET containing titanium dioxide (PET + TiO2), poly(methyl methacrylate) (PMMA) and poly(tetrafluoroethylene) (PTFE) has been used to enhance their wettability—which might be the first step in the immobilization of selected biological macromolecules such as heparin, and reduce interfacial tension between the biomaterial and a biofluid. The technique is still in the exploratory stage, the effects of exotic plasmas such as acetylene and ammonia on biocompatibility being explored, for example.

6.2. SURFACE ENGINEERING

6.2.2

105

Ion implantation

Ion implantation is a surface modification process in which positively charged high-energy ions, typically 10–200 keV, are implanted into the near-surface region of a substrate. The ions arrive at the target surface with kinetic energies 4–5 orders of magnitude higher than the binding energy of the host solid, and essentially form an alloy in the near-surface region. Commonly used ions are nitrogen and boron. To form the beam, an appropriate gas is fed into an ion source, in which electrons emitted from a hot filament ionize the gas to form plasma. The presence of an electrostatic field results in the acceleration of the positive ions at high energies under high vacuum (pressures below 10−5 Torr). The ions penetrate the target surface, typically down to 100 nm. Products of an ion implantation process typically include nitrides, borides or carbides. Ion implantation offers numerous advantages for treating component surfaces. The primary benefit is the ability to selectively modify the surface without detrimentally affecting bulk properties, largely because the process is carried out at low substrate temperatures. The process is also extremely controllable and reproducible and can be tailored to modify different surfaces in desired ways. Although it is a line-of-sight process, specialized fixturing can be used to uniformly treat complex geometries. Many functional properties can be improved with ion implantation, including hardness and wear resistance, resistance to chemical attack, and diminished coefficients of friction. Examples of components treated with ion implantation are Ti and Co-Cr orthopaedic prostheses, which are made harder and more wear resistant with the process, and silicone rubber catheters, which are made less tacky and more hydrophilic, improving insertion ability and biological compatibility (antithrombogenecity). There is evidence that ion implantation improves the lifetime of prostheses such as hip joints and knee components. Changing the local chemistry at the surface of the prosthesis may also result in beneficial effects, e.g. C+ and CO+ ion implantation in titanium and Ti-6Al-4V form dental implants result in better bone integration due to Ti-O-C covalent bonds at the bone-implant interface, stronger and more stable than the ionic bonds typically formed between TiO2 and bone. The treatment of polymers, e.g. by Ne+ or He+ bombardment, has also been shown to be beneficial, e.g. form better controlling cell adhesion on polymer surfaces. Ion implantation of polymers is routinely undertaken to improve the tribological properties of ultra-high-molecular-weight polyethylene (UHMWPE) used extensively in hip and knee implants (UHMWPE wear debris can initiate tissue inflammation, bone loss (osteolysis) and im-

106CHAPTER 6. SURFACE GENERATION, ENGINEERING AND COATING plant loosening).

6.2.3

Anodizing

A thin layer of TiO2 , typically 2–5 nm thick, forms naturally on titanium in oxidizing environments, generating a passive layer that provides excellent corrosion resistance. This oxide layer can be modified by anodizing treatments to increase its thickness and alter its morphology to enhance surface biocompatibility for orthopaedic implants. Typically the anodizing process consists of alkaline cleaning, acid activation and electrolytic anodizing to produce an oxide that may be between 20 and 1000 nm thick, with the thicker films exhibiting porous outer regions. The acid pretreatment removes surface contaminants and the general appearance of the anodic film gives a good indication of the homogeneity of an implant surface (e.g. Figure 6.9). Sul Figure 6.9: Scanning electron micrographs of titanium anodized in acetic acid showing two different types of surface morphology: (a) a non-porous barrier type in the control implant; and (b) a porous type in the test implant. (c) shows the honeycomb appearance of the porous implant at tenfold higher magnification (the scale bar in the top left-hand corner represents 10 µm). After Sul, Y.-T. (2002). The porous surfaces exhibit better histomorphometrical parameters (bone-implant contact and quality of newly formed bone) and osteoconductivity was more pronounced around the implants.

6.3

Surface coatings

This section deals with the deposition of a film of material on bulk objects in order to modify their surfaces. Topics covered include physical vapour deposition, chemical vapour deposition and thermal spraying technology. Attention will be paid to the control of deposit microstructure and coating composition via deposition parameters. Both thick and thin coatings are covered. The selection of a particular surface engineering route depends on many factors, including the substrate material, component design and geometry, cost and, obviously, the end application. Two aspects of the surface engineering process that are often highlighted are coating thickness and process temperature. The depths of surface treatment and thicknesses of coatings

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can vary over several orders of magnitude from less than 100 nm for ion implantation to several mm for thermal spraying and weld overlays. The importance of coating deposition temperature depends on the substrate material. The temperatures commonly encountered for surface engineering processes range from ambient up to the melting point of many substrate materials. High processing temperatures may adversely affect substrate properties but may be necessary to achieve the required coating stoichiometry and structure. This may be the limiting factor when selecting a coating and the deposition process.

6.3.1

Physical vapour deposition (PVD)

PVD processes include evaporation and sputtering. As with CVD (see below), there are benefits to using plasma-assisted deposition (PAPVD), in which the substrate is subjected to a flux of high energy ions both before and during deposition. The coating characteristics are a function of composition, microstructure and the deposition conditions. Simultaneous evaporation or sputtering from source materials of varied composition enables functionally graded or nanolaminated coatings to be produced. Also, temperature and pressure influence the nucleation and growth processes occurring, enabling control over coating microstructure and surface topography. PVD and PAPVD coatings have been applied to a range of biomedical applications. Much of the interest is in corrosion and wear resistance and the majority of work reported in the literature has concentrated on thin hard coatings. For example, TiN and (Ti6Al4V)N thin films deposited on to both silicon and steel by reactive magnetron sputtering exhibit excellent mechanical properties and corrosion resistance in prosthetics. Bioimplants such as hip joints and bone plates are prone to undergo fretting fatigue failures during service within the body. Thin hard coatings have been shown to improve fretting fatigue resistance because of superior tribological properties compared with uncoated alloys. PVD TiN, for example, has been shown to perform well on titanium alloys used in hip joints. Diamond-like carbon (DLC) coatings are popular for biomedical applications; titanium-alloyed DLC coatings have been deposited on Co-Cr-Mo substrates by PAPVD (magnetron sputtering) by using titanium targets in a mixture of argon and acetylene. Acetylene flow rate can be varied to generate a graded coating (titanium-rich at the substrate and DLC at the surface). Such coatings exhibit excellent mechanical properties and adhesion, low friction and exceptional wear resistance, and a biocompatible. DLC can also be applied to polymers. Cell attachment studies showed low macrophage attachment on DLC, which provided conditions for normal growth of fibroblasts.

108CHAPTER 6. SURFACE GENERATION, ENGINEERING AND COATING The number of neutral granulocytes and platelets adhering to DLC was also shown to be low, indicating that the coatings were highly biocompatible. The antimicrobial properties of implant surfaces are critical in preventing biofilm formation and deep infection of endoprostheses, and surface coatings can play a major role in modifying surface characteristics to minimize the attachment of microbes. Titanium/silver coatings, exploiting the cytotoxicity of silver and deposited by combined titanium evaporation and silver sputtering, appear to be effective. The silver is slowly released into the aqueous environment.

6.3.2

Chemical vapour deposition (CVD)

CVD and related deposition processes can typically produce coatings in the thickness range 1–100 µm. For the coating of biomedical components the desirable range is thin (1–10 µm) and based on ceramics that will provide enhanced wear and corrosion resistance. There are many variants of CVD, but essentially a workpiece is heated in a reactor into which a mix of reactive gases is introduced. Near to or on the workpiece surface a chemical reaction takes place to form a solid reaction product, which deposits as a coating. The chemical reactions may require high deposition temperatures, possibly resulting in problems with the thermal stability of the workpiece material. Therefore, in recent years there has been considerable effort aimed at reducing the temperature of CVD processes. Rather than thermal reactions only, plasma-assisted (or -enhanced) deposition (PACVD or PECVD) has proved to be very successful. Plasma-assisted reactions can occur at a the surface before and during coating deposition can influence both the adhesion and the structure of the deposit. There is considerable current interest in the use of diamond and diamondlike coatings for biomedical applications. They can be produced by PACVD as anti-infection coatings on implant surfaces. Medical implants have also been coated with diamond-like carbon (DLC) to increase wear resistance, reduce friction and provide corrosion protection. Because the DLCs are deposited by PACVD, they can be deposited on surfaces at low temperatures, which means that organic polymers, as well as metals and ceramics, can be coated. The DLC coating microstructure has been shown to possibly contain a diamond phase, a graphite phase and other allotropic forms of carbon. Crystal size was typically in the nanometre range and provide the coating with its excellent mechanical properties, a high strength passive film and good wear resistance. Histopathological investigation have shown that implants coated with DLC have excellent biotolerance. The coating also effectively protect against corrosion and metal-induced toxicity.

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DLC coatings have also been applied to the NiTi shape-memory alloys used for implants and orthodontic wires in order to improve biocompatibility. Patients may be allergic to nickel, but the PACVD DLC coating acts as an effective barrier that is essentially inert with respect to the local body environment. The benefits of the low temperature deposition afforded by plasma-assisted CVD have been demonstrated, controlling deposition conditions to produce thin films of amorphous hydrogenated carbon (C:H) coatings on polystyrene using a methane-hydrogen gas mix. The coated surfaces are not toxic to living cells, appear to increase cell attachment and afford normal cell growth rates. However, due to the low deposition temperature, the position of the substrate in the plasma is typically critical to ensure good coating adhesion. Hydrogen-free amorphous diamond coatings have the potential to enhance the wear resistance of total hip prostheses. Amorphous diamond films are biologically inert and highly adherent. Simulator testing of conventional total hip prostheses with either the ball or both surfaces coated demonstrate no detectable wear even after 15 million test cycles, equivalent to 15 years of clinical use. DLC coatings have also been applied to metal stent surfaces in order to reduce thrombogenecity by increasing the threshold at which solid foreign bodies trigger the clotting cascade. A key requirement for the coating is uniform and complete coverage to ensure that the generation of nickel ions is minimized. Small defects in the coating will result in crevice corrosion and localized dissolution of the metal substrate. The DLC coatings have been shown to be effective in preventing nickel dissolution from 316L stainless steel stents A further advantage of DLC coating is the ability to generate ultrasmooth surfaces, which are, as already discussed, critical for components such as stents. Recent studies have suggested mechanisms for the rapid smoothing that can occur even when a rough substrate is used, because of the erosion of hills into neighbouring hollows. Vapour phase deposition is also appropriate for the manufacture of highly porous structures. An example is “trabecular metal”, a highly porous biomaterial that is being developed for various orthopaedic application. It has an appearance similar to cancellous (trabecular) bone, having an open porous structure and a finely textured surface. The material is manufactured by pyrolysis of polyurethane foam to form a low density reticulated vitreous carbon skeleton that is then coated with tantalum by CVD or chemical vapour infiltration. The resulting material has up to 80% porosity and pore sizes typically between 400 and 600 µm (Figure 6.10). The interconnected pores promote greater bone ingrowth and the textured tantalum surface is highly osteophilic, promoting rapid bone growth.

110CHAPTER 6. SURFACE GENERATION, ENGINEERING AND COATING trab Figure 6.10: Trabecular metal (courtesy of Zimmer). See text for more details.

6.3.3

Ceramic coatings by plasma spraying

Plasma spraying is a material deposition process that belongs to the family of thick overlay coatings produced by heating a material in a hot gaseous medium and accelerating it at high velocity onto a substrate surface. Different types of thermal spraying process use various heating methods. In plasma spraying a DC electric arc is used to generate a stream of high temperature plasma, which acts as the spraying heat source. The coating material is normally in powder form and is supplied to the plasma in an inert gas stream where it is heated and accelerated towards the workpiece. The high temperature and high thermal energy of the plasma jet mean that materials with high melting points can be deposited. Hydroxyapatite (HAp), Ca10 (PO4 )6 (OH)2 , is a commonly used biomaterial for various types of prostheses encompassing orthopaedic, maxillofacial and dental applications. HAp is chemically similar to the mineral component of bone and may therefore be considered to be bioactive. It is applied to metallic implants of titanium alloys and stainless steels, most usually by plasma spraying. An example of a HAp coated hip prosthesis shaft is shown in Figure 12. The HAp coating is used to modify the surface characteristics in order to promote bone ingrowth. Key issues are the adhesion of the coating to the metallic substrate and the control of coating composition and microstructure during deposition and when in service. HAp is thermodynamically unstable at the high temperatures used in plasma spraying and this promotes the formation of CaO, which reacts with water and has a high solubility in body fluids. High deposition processes are also responsible for the formation of amorphous phases that reduce the coating-metal interfacial strength. Therefore recent developments have concentrated on improving coating stability and adhesion. Controlled atmosphere plasma spraying (CAPS) is useful to control the degree of melting of the HAp powder, enabling coatings of tailored microstructure to be produced. Composite yttria-stabilized zirconia incorporated into HAp contain more unmelted particles and greater porosity. The ZrO2 reacts with CaO to form CaZrO3 and bond strength increases. Microstructural control can be achieved by thermal treatment of the coating after deposition. Post-heat treatment promotes the complete crystallization

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of the amorphous calcium phosphate and this results in much improved adhesion. The poor bond strength that may arise between HAp and the metal substrate may be due to the mismatch in physical and chemical properties between the ceramic coating and metal. There is a large difference in thermal expansion coefficients, and rapid cooling of the plasma sprayed droplets results in residual stress in the coating which may compromise adhesion. Computerized closed-loop powder feed systems are now used to plasma spray functionally graded coatings by altering the ratio of HAp, Ti, Al2 O3 , and ZrO2 during the deposition. A typical microstructure coating is shown in Figure 6.11. It consists of a mixture of hydroxyapatite, bioglass (Al2 O3 , P2 O5 ) and zirconia. The composition changes continually so that the modulus increases to that of the titanium alloy substrate as the interface is approached. The outer layer is 30–50 µm of porous hydroxyapatite. Coating performance and long term stability are much improved through a reduction in tensile residual stress (from around 70 down to 10–2- MPa) and an increase in adhesive strength (from 20 to 50 MPa). Ning Figure 6.11: Graded coating. See text for more details.

6.3.4

Self-assembled coatings

This technology has the potential to be applied at low cost to surfaces of arbitrary shape to yield great chemical sophistication and variety. In the case of coating with polymers, once the polymer has been synthesized, in principle it can be applied simply by dipping the object to be modified in a solution of the polymer. The most common approach to polymer architecture is to create block copolymers, in which one block is made from a polymer that interacts strongly with the material of the object to be coated, and the other block has the required interactions with the environment. One of the most common applications of this principle is the use of poly(ethylene glycol) (PEG), i.e. poly(ethylene oxide) (PEO), as an agent to inhibit the adsorption of proteins. PEG (PEO) is the most hydrophilic substance known to man, which accounts for its remarkable ability to repel (hydrophilic) proteins, due to the very strong Lewis acid/base repulsive interaction. It should perhaps be mentioned here that many papers refer vaguely to “steric stabilization” as the reason for the repulsion of soluble proteins, but this ill-defined concept is unnecessary to explain the observed data, which are in quantitative

112CHAPTER 6. SURFACE GENERATION, ENGINEERING AND COATING agreement with the predictions of Lewis acid/base repulsion. A popular architecture is a block copolymer of PEG and poly(propylene glycol) (PPG), i.e. poly(propylene oxide), typically the triblock PEG-PPG-PEG. These materials are produced commercially on a considerable scale by BASF under the name Pluronics. The rather hydrophobic PPG section adsorbs strongly onto surfaces that would otherwise interact with, and probably denature, proteins, leaving the hydrophilic sections in the aqueous environment of the surface. There are endless variations (and patents pending or granted) on this theme. For example, for enhancing protein repulsion from titania, which is weakly negatively charged under physiological conditions, instead of PPG a polycation such as polyallylamine or polylysine will bind more strongly to the metal oxide. One disadvantage of this approach is however that any sudden change in the environment (e.g. lowering the pH) might lead to complete removal of the protective coating, with potentially disastrous consequences in a medical situation. Furthermore, the adsorbed polymer system is actually quite complex and still a very active field of theoretical investigation, and experimental results are by no means fully understood. At present, therefore, this technology has only found very limited applications in medicine. There is more promise in applications such as the prevention of biofouling on boats, for which somewhat different criteria apply, although the general principle is the same. Self-assembly of mineral particles is an equally powerful and convenient technique for modifying characteristics of a surface. Alternating polyelectrolyte deposition The method of alternating polyelectrolyte deposition appears to have immense potential. It requires the substrate to be electrostatically charged when immersed in water. It is then dipped into a solution of a polyelectrolyte of opposite charge, the excess is then washed off, and the coated substrate is dipped into a polyelectrolyte of the opposite charge, whose excess is again washed off, and so on (Fig. 6.12). polyelec Figure 6.12: Sketch of the polyelectrolyte deposition procedure.

There are few restrictions on the choices of polyelectrolytes. Early work was done with polyallylamine as the polycation, and polystyrene sulfonate

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113

as the polyanion. The essential feature of the technique is that at each dipping stage, the substrate charge is not only neutralized, but reversed, hence allowing the deposition to be repeated indefinitely. The exact role of the counterions, which of course are inevitably present,has not yet been elucidated. Each deposition is effectively irreversible: desorption would require the simultaneous release of hundreds of monomer units, which even if the probability for individual release is very high, rapidly becomes vanishingly small for n (the number of monomer units to be released simultaneously) greater then a few dozen (for example, 0.98400 ≈ 3 × 10−4 ). However, in sufficient excess of small counterions, it should be possible to desorb the polyions by a zipper mechanism. It has been shown that the deposition of each layer is esentially instantaneous, i.e. transport limited, and that where polymeric polyions are used, the layers strongly interpenetrate, such that the preponderant molecular orientation is perpendicular to the plane of the substrate, rather than laminar. Typically, polymer polyelectrolyte multilayers are labile at room temperature, as manifested in various ways; for example, it is not possible to produce images of the surface topography with an atomic force microscope—the chains are too flexible. Instead of polymeric polyions, particles composed of materials with ionizable surface groups (??) can be used. In this case, although the electrostatic charge of the surface of the coating is always reversed, not all the inner charges are compensated, because of steric hindrances, and hence electrostatic charges build up, and because the electrostatic force is long range, ultimately the coating as a whole becomes neutral and no further particles can be deposited. Another variant is to use block copolymers instead of simple polymers. This adds another degree of freedom to coating design. A very powerful extension of the technique is to use polymeric polyions as the polyelectrolyte of one sign, and ionizable particles as the polyelectrolyte of the opposite sign. In this was very sophisticated coatings can be built up. It has been established that the shells of many marine organisms, such as the abalone, are assembled using this principle, producing materials that are both beautiful and robust: anisotropic nanoparticles are dispersed in a biopolymer matrix, which only occupies a few volume percent of the total mass. Natural biopolymers, which are nearly all heteropolymers, primarily based on amino acids as monomers, but also possibly incorporating polysaccharides and nucleic acids, can incorporate enormous functional flexibility, in ways that we can only dream about at present in synthetic systems. The multilayer deposition technique should also be applicable to materials interacting via the da (electron donor-acceptor or Lewis acid-base) interaction (Chapter 2). In an important extension to this nanolamination fabrication technology, it has

114CHAPTER 6. SURFACE GENERATION, ENGINEERING AND COATING been found that the porosity of the layers can be reversibly controlled by varying the pH. It should however be pointed out that the biocompatibility of this composites is largely uninvestigated. abalone Figure 6.13: The nanoscopic structure of seashells. Dense, hard inorganic particles are bonded together by bioorganic polymers. The volume fraction of the latter is typically only a few percent. It is believed that the particles and the polymer are bound by opposite electrostatic charges.

The coming era of nanotechnology is likely to bring many new challenges to the province of fabrication of biomedical devices. The biocompatibility requirements of for example a “nanobot” have as yet been scarcely subjected to scrutiny. “Ready-to-use” integrated manufacturing procedures are likely to become more and more important.

6.4

Further reading

Allen, D.M., 2004, Photochemical machining: from manufacturings best kept secret to a USD 6 billion per annum, rapid manufacturing process, Annals CIRP, 53/2, 559572. M.F. Ashby, Materials Selection. De Chiffre, L., Kunzmann, H., Peggs G.N. and Lucca, D.A., 2003, Surfaces in precision engineering, microengineering and nanotechnology, Annals CIRP, 52/2, 561577. Evans, C.J. and Bryan, J.B., 1999, Structure’, Textured’, or Engineered’ surfaces, Annals CIRP, 48/2, 541556. A.G. Mamalis et al., Micro and nanoprocessing techniques and applications, Nanotechnology Perceptions, 1 (2005) 31–52. Pierson, H.O., 1999, Handbook of Chemical Vapor DepositionPrinciples, Technology and Applications, 2nd edn, William Andrew. Teoh, S.H., 2004, Engineering Materials for Biomedical Applications, Singapore: World Scientific. Rickerby, D.S. and Matthews, A. (eds), 1991, Advanced Surface Coatings, a Handbook of Surface Engineering, London: Blackie. Ramsden, J.J. and Mt, M., 1998, Kinetics of monolayer particle deposition, J. Chem. Soc., Faraday Trans., 94, 783788.

Chapter 7 The metrology of biomedical surfaces The metrologist concerned with biomedical surfaces in confronted with two tasks: firstly, to ensure that from the engineering viewpoint manufactured surfaces have the specified attributes (this will of course include characterizing the results of novel manufacturing procedures); and secondly, to investigate the response of a surface to its biological environment. The former task is of course common to all surface metrology, and rather standard techniques may in principle be used. Nevertheless, the variety of materials used in biomedical applications, and a variety of the attributes believed to be important for biocompatibility that therefore have to be quantified, make the task unusually challenging. A particularly difficult challenge is the apparent need to make measurements over an exceptionally wide range of length of scales: objects encountered in biological milieux range from the proton through relatively small organic molecules with relative molecular weights of a few hundred, macromolecules with molecular weights of one to several hundred thousand, and finally cells, ranging from bacteria that might only be a few hundred nanometres in diameter up to human body cells tens of micrometres in diameter. Many investigations are carried out with respect to a specific device, for which only a subset of the complete set of surface attributes may be relevant. For example, replacement joint implants require low friction rubbing surfaces, resistance to wear, and possibly the ability to retain lubricant. Stents for insertion into blood vessels need to be morphologically smooth, to avoid tearing tissue while being inserted by the surgeon, but once in place, the ability to resist protein adsorption may be very important. Microfluidic channels for biomedical devices merely need to have surfaces that do not impede the fluid flow, and so on. Hence, the actual metrological task may 115

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be somewhat simpler than might appear at first sight. The second task, essentially to quantify biocompatibility, is uniquely relevant to biomedical surfaces. It involves above all the investigation of the adsorption of proteins to the surface. Whereas academic investigations may focus on the adsorption of a single protein, or perhaps a simple mixtures of two or three different kinds of molecule, biological environments are vastly more complicated, and except in unusually favourable cases in which highly simplified systems can give useful information about the behaviour of implants in real life environment, the response of the surface to the real, complicated medium must be investigated. Hence, actual measurements of protein adsorption are almost indispensable. At the same time, efforts to correlate adsorption propensity with physico-chemical attributes of the surface have been made for many decades. If these efforts are ultimately successful, biocompatibility testing will become much simpler. At present, however, it remains difficult to predict the ultimate behaviour of proteins at surfaces, even when one is in possession of rather complete data regarding the surface morphology and chemistry. Furthermore, it is typically insufficient to merely record the accumulation of proteins at the surface. Very often the surface becomes coated with a protein monolayer, which remains quite stable: such a surface would be considered to be biocompatible if the protein retained its native conformation. In most cases, however, significant protein conformational changes take place, and as a result the protein is able to trigger extraordinary interactions with the body, e.g. inflammation, innate or even adaptive immune response. According to the arguments developed in Chapter 1, it is above all the surface-protein interfacial energy that determines whether a protein will be able to arrive at, let alone adhere to, a surface. This may even be considered to be a mesoscopic parameter, which in turn can be correlated with the microscopic (i.e. nanoscale) morphological and chemical features. For all those applications in which the surface is that of an implant destined for long-term residence in the human body, the ultimate goal is assimilation by the host tissue, and therefore biomedical surface metrology may well have to include appropriate cytometry in order to characterize how cells interact with the surface. Since a full characterization of the cell involves thousands of parameters, such as the abundance of all the proteins it contains, this could become a very complicated affair indeed. Again, knowledge that permits a strongly restricted subset of parameters to be taken as representative of the cell in its entirety is at a great premium. The environment of the biomedical surface metrologist is very different from ultrapure surfaces in ultrahigh vacuum chambers. Many relevant measurements take place at the solid/liquid interface, which is strictly speaking

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a buried interface. In such an environment contamination is much more difficult to avoid than in the vacuum chamber. Even the most carefully purified liquid constitutes a considerable potential reservoir of impurities. On the other hand it might be argued that biocompatibility, even in its molecular interpretation in terms of protein adsorption as a fundamental process, does not depend on the presence of adventitious impurities, not least because in nature they may be thousands of different atomic and molecular species present, at concentrations ranging from a few to molar. This perhaps a reasonable heuristic viewpoint commensurate with the present state of knowledge, but it has not been definitively established. Because of the multifarious nature of many biomedical surfaces, great value is placed on multifunctional instruments that can simultaneously measure a number of distinct parameters. Indeed, it might well be asserted that they need to characterize biomedical surfaces has been the prime driver of the development of such multifunctional instruments. An example of such an instrument would be a stylus for measuring topography that can also detect the chemical nature of the zones over which it passes; an example of an actual instrument with this capability is a scanning electron microscope equipped with energy dispersive X-ray spectroscopy, or a transmission electron microscope capable of carrying out selected area electron diffraction (in this example it is the symmetry of the atomic arrangement that is simultaneously measured alongside some geometrical attributes of the shape of material, rather than the chemical identity of the constituent atoms). In the sections that follow, the arrangement is the same as in Chapter 2 and.

7.1

Energetics

Most of the arguments developed at the beginning of the book regarding protein adsorption, seen as the fundamental process underpinning biocompatibility or bio-incompatibility, were cast in terms of energy (i.e. the interfacial energy between adsorbate and adsorbent). Other parameters will have different effects. Roughness, for example, will affect the capacity of the surface (i.e. the total number of particles with which it can be covered by a monolayer) in the first instance; the mean curvature, it is significantly different from the infinite curvature of a planar surface, were affected the less attention, in the second instance. The most straightforward measure of surface energy is the ability of different liquids to wet the surface (cf. §).

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7.1.1

Wetting

The individual terms of eqn (1) can be calculated from contact angle data. In the absence of electrostatic charges, each term has three components, the Lifshitz-van der Waals potential, the electron-donating potential (dativit`a) and the electron-accepting potential (recettivit`a). Hence the contact angles of three different liquids (typically water, an apolar liquid such as bromonaphthalene, and a third liquid such as formamide whose dativit`a and recettivit`a are different from those of water), with the material under investigation must be measured, and the three corresponding Young-Dupr´e equations: (LW)

(LW) (LW) 1/2 γ2 )

+ (γ1⊕ γ2 )1/2 + (γ1 γ2⊕ )1/2 (7.1) solved simultaneously to yield the three component potentials of the material, γ (LW) , γ ⊕ and γ . Simple and convenient as contact angle measurements are in principle, there are in practice numerous practical pitfalls, and the number of laboratories in the world actually capable of producing reliable measurements is far smaller than the number possessing equipment with which to carry out the measurements. In fact, practically no equipment other then a protractor for measuring the angle is required. Rather successful results have been obtained with no more than a simple zoom camera with which the droplet is photographed. One assumes that the surface on which the droplet is placed is in the relevant state: i.e. if the surface energy of a “dirty” surface (e.g. one after having been exposed to a biofluid) is to be determined, obviously any cleaning is inappropriate; on the hand if the surface energy of the manufactured surface of an implant is to be determined, that would undergo exhaustive cleaning before been surgically inserted into the human body, that same cleaning procedure must be applied to the surface before the contact angle measurement. Most materials interact with that very powerful solvent and reagent, water. Hence especially for measurements of the contact angle of water with the surface, care must be taken to ensure that a pristine surface is exposed to the liquid, i.e. it must not be already hydrated (the measurement of the hydrated surface gives additional information, which may also be useful). The easiest way to ensure the pristine nature of the surface is to measure the advancing contact angle: i.e. liquid droplets must be set in motion such that the material beneath the advancing edge never has time to become hydrated. For obvious practical reasons, the notion should be as slow as possible consistent with achieving that aim. An easy way in of achieving such constant motion is to tilt the surface, whereupon the droplet will slowly and steadily move under the influence of gravity. The trailing edge of the droplet gives [γ2

/2 + (γ2⊕ γ2 )1/2 ](1 + cos θ) = (γ1

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the receding contact angle, but its value evidently depends on the contact duration between the liquid on the substance, which in turn depends on the size of the droplet and the speed of motion. Other important factors include the droplet size. A suitable series of measurements should be able to establish experimental conditions under which the intrinsic contact angle can be determined with reasonable accuracy. Table 7.1.1 gives the single substance surface tensions for a variety of useful liquids. Given some of the practical difficulties of obtaining precise and accurate values of contact angles, it is not surprising that other methods for measuring wetting parameters have been developed. One useful one is wicking,1 defined as the measurement of the contact angle θ that a liquid makes with a solid by determining the rate of capillary rise of the liquid in packed beds of the solids, using Washburn’s equation h = (tRγL cos θ/(2η))1/2

(7.2)

where h is the height to which the liquid has risen in time t, R is the effective interstitial pore radius between the particles, γL is the surface tension of the liquid, and η is its viscosity. γ (LW)

γ⊕

γ

water

27

6.3

51

α-bromonaphthalene

37

0.1

38

formamide

32

4.5

56

glycerol

32

4.5

56

substance

η/mN s m B.P./◦ C 1.28

100

Table 7.1: Surface tensions/(mJ m−2 ) of a selection of liquids useful for surface tension measurements. Viscosity (η) and boiling point are given because of their practical importance in choosing liquids for the experimental measurement of contact angles.

7.1.2

Probe particles

As Spielman and Friedlander have shown,2 the particle adsorption rate coefficient ka bears a simple relationship to the particle-adsorbate interfacial 1

C.J. van Oss et al., Determination of contact angles and pore sizes of porous media by column and thin layer wicking, J. Adhesion Sci. Technol. 6 (1992) 413–428. 2 L.A. Spielman and S.K. Friedlander, Rˆole of the electrical double layer in particle deposition by convective diffusion. J. Colloid Interface Sci. 46 (1974) 22–31.

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energy: ka = D/δa

(7.3)

where D is the protein diffusion coefficient and δa is the characteristic adsorption interaction length, which is linked to the interfacial energy (∆G(total) = ∆G(LW) + ∆G(da) + ∆G(el) ; cf. Chapter 1) according to Z ∞ δa = [exp ∆G(total) − 1]dz . (7.4) 0

Hence, the particle-adsorbate interfacial energy can be determined provided the adsorption rate coefficient can be determined with reasonably good accuracy (the exponential dependence in equation (7.4) then ensures that the interfacial energy can be determined with very good accuracy). Experimentally, this implies that one needs a system of counting the accumulation of particles on the surface very accurately. The two approaches to determining the adsorption rate coefficient are: 1. measure the initial adsorption rate as accurately as possible, and extrapolating back to zero time when the adsorbate is totally empty; 2. measure the complete adsorption kinetics, and use a reliable expression for the available area function φ The rationale for the two approaches is in both cases the canonical adsorption equation: dν/dt = ka c∗ φ(ν) (7.5) where ν is the number of adsorbed particles per unit area and c∗ is the concentration of particles in the immediate vicinity of the surface. Reasonable approximations for c∗ have already been discussed (Chapter 1); approach (1) assumes that φ = 1, and approach (2) can only work in the case of spheres or spherocylinders of known dimensions, hence allowing an analytical expression for φ to be used in equation (7.5).3 In the formal approach, a simple graphical method may be adequate to yield the desired rate coefficient; in the latter approach, equation (7.5) would normally be fitted to the experimental data in order to determine ka .

7.2

Morphology

As before, this comprises the surface topography and porosity. 3

See for example P. Schaaf, J.-C. Vogel and B. Senger, Irreversible deposition/adsorption processes on solid surfaces, Ann. Physique 23 (1998) 1–89, for a comprehensive discussion of available area functions.

7.2. MORPHOLOGY

7.2.1

121

Contact techniques

The rationale for grouping several techniques together under this heading is that an actual object possessing (rest) mass and extension is brought into contact with the surface under examination. The conventional methods are based on dragging a stylus over the surface, whose vertical motion (deflexions from a mean) more-or-less faithfully mimics the surface topography. The stylus being attached to a position transducer, a record of the topography can thereby be obtained. Obvious limitations are related to the dimensions of the stylus and the velocity of its motion parallel to the surface. An infinitely thin stylus moved with an infinitesimal velocity would be able to perfectly record every detail of the topography; in reality, finite thickness and finite velocity combine to make the recorded topography only an approximation of reality. Nevertheless, this “limitation” can be used to advantage: if the interaction with 10 micrometre micropebbles is the object of interest, a stylus a few micrometres in diameter will be adequate for obtaining a relevant recording of the surface topography, which will be obtainable much more rapidly then if the topography were measured down to nanometre resolution. As always, the key to success in metrology is the choice of an appropriate technique. Conventional methods The basic instrumentation comprises the probe stylus to which some pressure can be lightly applied, ensuring that it remains in contact with the surface, some means of moving it, in raster fashion, across the surface under investigation, and finally a method of recording the vertical motions of the stylus. This last essential could for example combine a transducer converting position into an electrical signal, such as an induction coil with a movable core, the electrical signal being recorded either digitally or on paper. Naturally a great deal of sophisticated detail goes into the design and construction of practical instruments, especially those now manufactured and sold commercially, requiring relatively little expertise in use. Contacting instruments with a large dynamic range, i.e. where the resolution of vertical height measurements enables the texture information to be extracted directly from the data, are useful when assessing the performance of, for example, hip joints and dental implants. Contacting techniques are especially advantageous where very steep-sided features need to be evaluated. The analogy with a gramophone needle will not have escaped attention. If large areas need to be evaluated statistically, the stylus could be scanned rapidly over the surface and its vibrations in the acoustic range decomposed into Fourier components, yielding a compact description.

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Scanning probe microscopies (SPM) Historically, the first microscope of this type to have been invented was the scanning tunnelling microscope (STM).4 Conceptually, the basic elements are just the same as those of the macroscopic stylus. The details are of course somewhat different: the stylus is a microscopic needle affixed to a flexible cantilever that acts as a spring (Figure 7.1). The sample is affixed to a piezoelectric crystal permitting its movement relative to the (fixed) needle in the x and y (i.e. the plane of the surface) and z (perpendicular to the surface) directions. The sample and the needle must be electrically conducting. They are brought sufficiently close to one another by applying voltage to the appear to make a crystal such that a tunnelling current I ∼ exp zmin

(7.6)

passes between the tip of the needle and the closest point on the surface, a separation of zmin . Keeping the z position fixed, the sample is moved rasterwise in the (x, y) plane while recording the tunnelling current. Due to the exponential dependence on separation (equation 7.6), the measured current is extremely sensitive to fine topographical details of the surface. Not many more years elapsed before the atomic force microscope (AFM) was invented in the same laboratory.5 It is an obvious extension that enables the scanning probe principle to be used with any kind of material, not only electrically conducting ones. The basic setup is practically identical, except that the physical deflexion on the cantilever caused by the Born repulsion (Figure 1.1) is recorded, either using the optical arrangement is sketched in Figure 7.1 or mounting a tunnelling probe above the atomic force cantilever. In an alternative mode of operation, the deflexion of the cantilever is kept fixed at a preset value via feedback between the deflexion transducer and the piezoelectric crystal controlling the deflexion. In principle the results should be the same as that of the direct deflexion measurement, but in practice they are not. In a third mode of measurement, the feedback activity is recorded, again with different results. In practice, the wise experimenter will investigate all the different possibilities and compare the results with those obtained with different techniques such as scanning electron microscopy— bearing in mind that they will each introduce specific distortions into the apparent measured topography. The exponential dependence of tunnelling current and distance means that the precise shape of the tip of a scanning tunnelling microscope needle 4 5

Bednorz Bednorz

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123

is not of great importance: only the atoms at the extremity will actually be involved. With the atomic force microscope, on the other hand, the tip shape strongly influences the measurement. Just as in the case of the macroscopic stylus, it is impossible to discern features at a scale finer than the characteristic dimension (e.g. the radius of curvature) of the tip of the needle. It is relatively straightforward to fabricate silicon or silicon nitride tips with a radius of a few tens of nanometres using standard procedures of semiconductor processing.6 As well as being cheap enough to be considered as disposable items, such tips are moreover fairly robust and would not usually break during measurement. Unfortunately many of the most interesting features that one would like to investigate with the atomic force microscope are smaller than 30–40 nanometres. Since the tip of the microscope is not infinitely sharp, the apparent lateral dimensions of objects are broadened. If r is the true object radius, the apparent lateral dimension L of an object imaged by a tip of radius R is given by L = 4(Rr)1/2 .

(7.7)

Some improvement can be achieved by independently measuring the precise shape of the tip using a scanning electron microscope and then processing the recorded deflexion profiles in order to deconvolute the influence of tip shape—this is essentially the modulation transfer function (MTF) approach familiar in optical image processing. A survey of current and past literature shows that, regrettably, this appears to be done rather infrequently. The alternative is to fabricate much finer tips. Until recently, this was done by controlled etching of the standard tips. Not only are these ultrafine tips mechanically much more fragile and hence easily broken during scanning, they are also laborious to make (if one does it oneself) or expensive to purchase. Furthermore, the fine tip is very easily contaminated by macromolecules and other debris. This is a particular problem if a surface coated with proteins is being investigated. The atomic force microscope (Fig. 7.1) has become a key instrument of nanoscience and nanotechnology and is ideally suited to to the quantitative determination of parameters characterizing biomedical surface structures at the nanoscale. Apart from potentially excellent resolution, particularly perpendicular to the surface, it can be used in ambient conditions, e.g. in air or in liquids. For imaging the biomedical surfaces after they have been in contact with the human body this is, indeed, almost indispensable: one of the permanent criticisms of electron microscopy is the need for considerable sample preparation, 6

See A.G. Mamalis et al., Micro and nanoprocessing techniques and applications, Nanotechnology Perceptions 1 (2005) 31–52, for a highly readable account of these procedures.

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including dehydration, before imaging can be carried out. Note, however, that the fairly new environmental scanning electron microscope (ESEM) enables images to be obtained under much less extreme conditions: a few millibars of pressure are acceptable, which means that lower lifeforms such as insects and worms can even be imaged alive. AFM Figure 7.1: Diagram of an atomic force microscope (not to scale). The sample is fixed on the platform P, in relation to which the block B can be moved in the x, y and z directions. A flexible cantilever is mounted on the block, and on the end of the cantilever is a sharp tip. The beam from a diode laser L is directed onto the cantilever, from which it is reflected onto a split photodiode D. The block is brought up to the sample such that the tip is almost in contact with it, and then scanned across the sample, i.e. in the x, y plane. Sample declivities cause the tip to move up and down, and the corresponding deflexions of the cantilever are faithfully recorded via the ratio of the signals from the two halves of the photodiode. Sophisticated signal processing allows the three dimensional topography of the surface to be extracted. By using a more elaborate setup with a quadruple split photodiode, sideways deflections of the cantilever can also be recorded, enabling the friction coefficient of the sample surface to be quantified at the nanometer scale. The basic output from the atomic force microscope is an array of height values h for the surface considered as an array of discrete spatial cells (see Table 2.2 for roughness formulae). A wealth of additional information is however obtainable from the atomic force microscope. A simple extension to the instrumentation (see the legend to Figure 7.1) allows the friction coefficient of the material to be determined with the same resolution as for the topographical information. Since the microscope tip can be modified in as many ways as ingenuity can dream up, the friction of contact with any given material can be measured simply by fabricating an appropriate tip and mounting it on the microscope cantilever. Furthermore, tips made from different materials can be chosen in order to investigate the combined effects of interfacial forces and surface topography (defined by the hard body or Born repulsion). Careful work with the atomic force microscope should begin with measurement and analysis of the forcedistance characteristics (see the next section). There is no such thing as an

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125

absolute topography at the nanoscale, merely a topography in relation to the minimum repulsive force in effect. If a rather stiff cantilever is used, imaging may faithfully follow the topography of the Born repulsion. A more flexible cantilever may be sensitive to the longer-range electron donor-acceptor (hydrophilic) repulsion and may yield a quite different topography (unfortunately continuously variable cantilever stiffnesses are not yet available). Complete characterization of the surface—and this is especially important for biomedical applications, knowing that biomacromolecular interactions are strongly influenced by these “soft” forces—requires measurement with a range of tips, e.g. of increasing hydrophobicity. As already mentioned, an important advantage of scanning probe microscopy is that it can take place in natural media. Unfortunately, it can scarcely be recommended to attempt to image a biomedical surface in the presence of blood (serum) or other protein-containing biofluids such as tears, because the proteins will tend to deposit on the probe tip, continuously changing its shape. Hence a series of model measurements under different idealized conditions from which results applicable to the real biomedical environment can be extrapolated are still required. In practice, it may be very troublesome to procure tips made from different materials (corresponding to for example the range of hydrophobicities suggested above), especially since in order for the measurements to be comparable with one another, the tips would have to have identical shapes and sizes. Therefore, more effort has been put into modifying tips made from a standard material (such as silicon). Using such an approach, not only can very thin films (a few nanometres) of different inorganic materials be deposited using techniques such as physical vapour deposition (see Chapter 6), but also soft macromolecules can be grafted onto the tip, permitting for example the direct measurement of the adhesion forces between a cell surface integrin and a deposited extracellular matrix layer. Another fashion of extending the capabilities of scanning probe microscopy is to record not the tunnelling current that the faradaic (electrochemical) currents between a microelectrode as the scanning probe and a conducting substrate. The lateral resolution in this case is limited by the rates at which the ions to carry the current diffuse at random.

7.2.2

Porosity

The customary technique for measuring porosity is to consecutively admit known amounts of a gas that can be adsorbed by the pores into a chamber containing the sample of material was porosity is to be measured, which may be kept at low temperature to favour the adsorption (e.g. that of liquid nitro-

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gen, 77 K). The pressure then slowly decreases, and by the new equilibrium is reached the next aliquot of gas is admitted. By this means an isotherm of the amounts adsorbed (typically per unit mass of the sample) versus equilibrium concentration (pressure) can be obtained, and interpreted in terms of model. The simplest one is that of Langmuir, according to which the surface is covered with discrete noninteracting adsorption sites individually larger than those of the adsorbing probe gas molecule, yielding an isotherm saturating exponentially. A very widely used isotherm is a generalized Langmuir model alike and multilayers, the Brunauer-Emmet-Teller (BET) isotherm. Ideally isotherms should be measured for a range of gases similar chemically but of increasing size. Nevertheless, the practical range of such measurements is far lower than what would be required for example to establish a fractal dimension of the porosity. The pore size distribution is typically determined by fitting to the actual measured isotherm. Techniques based on gas adsorption can only measure the open pores. If it is desired to determine the closed pores, then a method such as X-ray tomography has to be used.

7.2.3

Non-contact (optical) techniques

Microscopy Measurement methods based on using the characteristics of light interacting with surfaces has been used extensively for many years. Their main advantage is that the surface under test is seldom physically damaged, or chemically changed, by their use. Optical profilers are analogous to the mechanical stylus instruments but use focused beams to detect the location of the surface.7 A very important addition to the scanning probe microscopy family is the scanning near-field optical microscope (SNOM). The scanning arrangements remain the same, but now an optical fibre brings light very close to the surface (see Fig. 7.2). Transmission, reflexion and fluorescence can all be measured. The obtainable resolution is below the diffraction limit applicable to far field optics. Such measurements open up new perspectives in the determination of optical surface properties. 7

Laser vibrometry is a dynamic mode of such measurements. See Rembe, C., Dr¨ abenstedt and Heimes, F. (2005). An accurate new 3D-motion analyser for MEMS and bioMEMS. J. Biol. Phys. Chem. 5, 37–40.

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127 SNOM

Figure 7.2: The principles of scanning near field optical microscopy (SNOM). Left hand diagram: SNOM in illumination mode: a dielectric probe illuminates the sample from above, the light is collected in the far field; dashed line: reflexion mode in which the radiation reflected off the surface is collected by far field optics (alternatively the probe tip itself could be used). Right-hand diagram: SNOM in collection mode: the sample is illuminated from below and a dielectric probe collects the light coming through the sample. The two critical requirements are: (i) a sub-wavelength light source, and (ii) placing the sample in the near-field zone of the light source.

Light scattering techniques Microscopy is based on images, which penetrate powerfully into the human psyche, although they may not be easy to interpret quantitatively, and their scientific value is often diminished because insufficient images are examined to determine whether those selected for detailed examination are truly representative. The alternative approach is directed towards statistical, nonimaging methods based on light scattering. Conventional techniques include specular reflexion, total integrated scattering, and angle-resolved scattering. Speckle techniques include speckle contrast, speckle pattern illumination, and angular- or wavelength-dependent speckle correlation. New technologies, including powerful light-emitting diodes and diode lasers, fast CMOS and CCD cameras, and, not least, powerful information processors, have moved these techniques into the domain of real time surface roughness measurement. The diameter of the zone over which the measurement is integrated typically ranges from 0.5–10 mm, and roughnesses from 1 nm to 10 µm are accessible. The speckle pattern illumination method, based on doubly scattered coherent light, is suitable for measuring the roughness of specularly reflecting surfaces with roughness ranging from 1 to about 125 nm, corresponding to the surfaces of hip replacements, stents and eye lens implants. Experimentally, the surface is illuminated with a monochromatic fully developed speckle pattern, whose phase distribution is modulated by the rough surface. Polychromatic speckle autocorrelation involves illuminating the surface with a collimated, partially coherent (i.e. polychromatic) light beam, either discrete (produced by a combination of laser diodes) or continuous (produced by superbright light-emitting diodes for example). The method is suitable

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for diffusely scattering surfaces with a root mean square roughness of less than a quarter of a wavelength, i.e. about 150 nm. The resulting scattered light intensity distribution shows a roughness-dependent speckle elongation (Figure 7.3). An efficient procedure to obtain a numerical parameter starts from the ratio of the mean speckle diameter Dx (estimated using a local two-dimensional autocorrelation function) in the horizontal direction to the mean speckle diameter Dy in the vertical direction, calculated for different segments of each speckle pattern. speckle Figure 7.3: Polychromatic speckle patterns for (a), upper panel, Ra = 0.4µm; and (b), lower panel, Ra = 6.4µm. In the former case, the speckles are elongated, and in the latter, completely decorrelated.

7.2.4

Layer thickness and density

The techniques discussed in this subsection are mainly applicable to the study of thin films (coatings) overlaying a bulk substrate. Neutron reflectivity has the advantage that the wavelength of the probe radiation (i.e. neutrons) has a wavelength of the same order of magnitude as the layer thickness. It is possible to characterize the density profile perpendicular to the plane of the film. Disadvantages include the need for a powerful neutron source, which means that it cannot be used routinely in the laboratory, and it generally required deuteration of the normally hydrogencontaining molecules constituting the film. Classical thin film optical techniques such as ellipsometry or scanning angle reflectometry can be used to determine the optical parameters (reflectance) of the surface of a solid object, and the thickness and optical parameters (refractive indices) of thin films coating the surface. Provided the instrumental design permits insertion of the object being investigated, they can even be regarded as being in situ techniques. These techniques are also very useful for investigating films of biomolecules deposited on an implant in service. Ellipsometry is most useful for determining the total quantity of material a unit area in the film, but scanning angle reflectivity is more versatile and can be used to obtain the structural information about adsorbed biofilms, including the density profile perpendicular to the plane of the film.8 8

E.K. Mann et al., Optical characterization of thin films: beyond the uniform layer

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For the development of novel coatings, in which an in situ technique is not required, it is advantageous to use an optical waveguide as a substrate. The spectrum of guided lightmodes yields precise information about the optogeometric parameters of the film—one parameter for each lightmode whose propagation constant is measured. Using the uniform thin-film approximation, layers of discrete particles of thickness not exceeding about one tenth of the probe wavelength can be modelled as uniform thin films with insignificant error.

7.3

Dynamical structure measurements

Viscoelasticity If material is allowed to accumulate on a resonantly vibrating object, its resonant frequency diminishes as mass is built up. The most familiar example of technology based on this principle is the quartz crystal microbalance (QCM). Quartz is a convenient piezoelectric material and by depositing electrodes on opposite faces of the crystal, it can be set in vibration. The QCM is well established in vacuum chambers for monitoring physical vapour deposition processes, but the data are much more difficult to interpret when used with viscoelastic biodeposits in the presence of water, although careful scrutiny of the dissipation seems to be a promising approach towards extracting less ambiguous and more interpretable results from the data.9 Surface acoustic wave (SAW) devices are a variant of this technique, and the vibrating cantilever is a typical MEMS realization that could form the basis of high throughput, massively parallel devices useful for combinatorial optimization of biomaterials. Although the output from such devices does not have the straightforward interpretation of the output from an OWLS experiment, large-scale screening might not require that level of detail and reliability. Non-imaging uses of scanning probe microscopy The atomic force microscope can also be used to determine interfacial forces directly. Typically this begins by imaging the sample in order to determine a favourable position for subsequent force determination. Then, at one selected point on the sample surface, the needle is slowly made to approach the sample (starting from position o—see Fig. 7.3, at which it is not in contact model, J. Chem. Phys. 105 (1996) 6082–6085. 9

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with the sample). At point a the detector response moves from a to a0 . This corresponds to movement of the monitoring beam reflected off the cantilever from an initial position near the perimeter of the four segment diode detector towards the center of the four segments, without the cantilever being deflected. During the displacement a → a0 the beam is centered on the detector and a practically constant response, indicating a practically constant deflexion, that is a practically constant force, is observed. This would correspond to tip movement in a viscous liquid such as a tangled fiber network. In this case Stokes’ law F = f ηv + F0 (7.8) should apply, where F is the applied force, η the viscosity of the liquid, v the velocity of tip movement, and f the characteristic dimensions of the tip. Plotting force against velocity allows the viscosity of the layer to be determined. Moving the tip still further into the sample, from b → c, causes a significant increase in detector response. The displacement z in this region is the sum of cantilever deflection dv and cellulose layer deformation dc , i.e. z = dv + dc

(7.9)

hence the response F can be written as F = kc dc − kv dv

(7.10)

where k is a parameter of deformability, provided Hooke’s law applies. Since kv /kc = dv /dc

(7.11)

kc2 − kv2 kc + kv

(7.12)

we obtain F =z

from which kc , the fiber layer deformability, can be obtained from the slope dF/dz. This type of approach is also used to study nano-indentation hardness. FDC Figure 7.4: Prototypical force-displacement cycle on the atomic force microscope. See text for full explanation.

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Tribology The Tabor-Eldredge instrument is still very important for friction measurements. The main part of instrument is a balance arm with a bearing at the fulcrum. The tangential force at one end of the arm is measured by a flexible steel beam attached to a strain gauge, and to which the probe is rigidly attached. Loads are applied to the probe by placing a calibrated weights above the probe. A counterweight is placed at the other end of the arm for balancing it horizontally prior to use. More elaborate instruments for measuring dynamic friction consists of the loadable rotating ball of all material placed on an independently rotating turntable from another material, on which lubricants can be dispersed.

7.4

Chemical surface analysis

In the same way as there are a host of optical and non-optical methods for analysing the physical attributes of a surface, there is an equally wide-ranging host of spectroscopic and other analytical techniques that enable information to be derived about the chemical condition of the surface. Techniques that can be used in the presence of moisture/water—highly relevant for biomedical surfaces—include Raman spectroscopy, but many of the techniques, such as ion spectroscopy and electron spectroscopy, require that the specimen under test be placed in high vacuum. All the techniques are really the same in principle: a beam of preferably monoenergetic photons hν is fired at the surface, and the reflected photons, bearing some traces of their interaction with the surface, are collected and characterized, primarily in terms of their energy hν ∗ ; in more specialized techniques the polarization and the phase may also be investigated. These reflexion-based techniques, one could say almost by definition, probe surface layers of the material under investigation, the depth of probing depending on the extinction coefficient of the material for the wavelength in use. The techniques can simply be classified in accordance with the photon energy. We begin with the lowest. In the infrared domain, the classic technique is infrared spectroscopy, which is particularly useful for investigating the presence of biological molecules, including proteins, present as a thin film on the surface of a different bulk material. Infrared photons are absorbed by interatomic bonds. A monochromatic the tuneable photon source is scanned through the relevant energy range and the intensity of reflected photons recorded for each energy hν. A dip in intensity at the particular frequency gives the energy of the bond.

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Raman spectroscopy falls into the same energy domain, the difference is simply that the quantum mechanical selection rules for allowed transitions are different. Important for biomedical surface investigations is the fact that ubiquitous water, infrared opaque, is Raman transparent. Raman-active adsorption bands can be enhanced by resonant interaction with a (visible) optical absorption band, corresponding to an electronic transition. This can be very useful when investigating a thin (submonolayer) film of e.g. an adsorbed protein, since several important metalloproteins have intense absorption bands in the visible spectral range. Without resonant enhancement, the signal-to-noise ratio from a submonolayer might be inadequate for measuring absorption. Further enhancement can occur with molecules adsorbed on noble metals such as silver (the so-called surface-enhanced Raman scattering, SERS), useful if such a metal happens to be the adsorbate. A multiphoton variant, also useful in the case of humid samples, is to direct a beam of visible light (hν1 ) together with a (tuneable) beam of monochromatic infrared light (hν2 ) onto the surface, and determine the intensity of the sum frequency h(ν1 + ν2 ): dips in intensity will occur when energy quanta (hν2 ) are absorbed by bond vibrations. This energy domain is also explored by some exotic spectroscopies using electrons as the incoming radiation, such as EELS (electron energy loss spectroscopy). In this case it is the difference in the energies rather than intensity is E0 − E1 = hν that is observed, the quantum hν corresponding to the bond energy. At slightly greater photon energies, one is in the realm of visible optical absorption spectroscopy, in which photons excite valence electrons of the constituent atoms of the surface under investigation to higher energy levels, but most of the compounds of interest to the biomedical surface engineer are transparent in that region. If the excited atomic state returns to its ground state with the emission of a photon, the phenomenon is called fluorescence, in which case it is useful to measure both the absorption and the fluorescence. The latter usually means the fluorescent emission spectrum; the fluorescence excitation spectrum may be compared with the absorption spectrum to gain insights into the nature of the material. In the X-ray domain, the incident X-ray photon of energy hν directly ejects an electron from the core level of an atom, which emerges with the kinetic energy given by Ekin = hν − Eb − eφ (7.13) where Eb is the binding energy of the core electron (i.e. the energy of the defect electron (hole) state produced by removal of the electron) and φ is the work function of electron spectrometer used to measure Ekin . Chemical

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bonds between the atoms in the solid affect the electronic charge distribution and therefore cause shifts in Eb compared with those characteristic of the free neutral atom. This is useful for determining the oxidation state of an element in the surface film, for example. The measurement of the spectrum of ejected electron energies is called X-ray photoelectron spectroscopy (XPS). Although X-rays can penetrate relatively deeply (maybe many micrometres) into most solids, only electrons produced near the surface (typically between one and 10 nm) have a significant probability of escape. The Auger effect, named after its discoverer, involves the ejection of an electron from a core level (as in XPS) by an incident electron beam. The vacancy is unstable and is filled by an electron from a high level, releasing energy that is either transferred to another (so-called Auger) electron (from the yet higher, i.e. more shallow, level), or emitted as an X-ray photon. The measurement of the spectrum of Auger electrons is called Auger electron spectroscopy (AES). The elastic scattering of X-rays is of course also employed to determine structures (i.e. interatomic spacings) of thin surface films. The final group of techniques to be discussed in this section concerns ion beams, typically Cs+ or oxygen ions. They may be used as highly energetic photons for imaging, rather in the manner of electron microscopy, but also secondary ions dislodged from the surface by the primary (incident) ion beam can be collected and mass-analysed in the spectrometer—this is called secondary ion mass spectrometry (SIMS). Ion bombardment is also used in combination with all the other techniques in order to erode the surface in highly controlled way and expose underlying layers not accessible otherwise.

7.4.1

Layer chemistry and multicomponent distribution

The methods described above offer no means of identifying the nature of the adsorbed matter, other than by inference from features of the adsorption kinetics, etc. This makes interpretation problematical if mixtures of proteins are being used. Antibodies have been used to identify the presence of particular proteins in an adsorbed layer deposited from a complex mixture such as blood plasma, but despite, or perhaps one should say because of, the exquisite sensitivity of antibodies to particular conformational motifs in the target protein, the method can only be regarded as semi-quantitative, because of the possibility that the protein sought for is present, but the relevant conformational motifs are inaccessible or sufficiently altered due to

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the assembly of protein in the adsorbed layer such that they are no longer capable of binding to the antibody. Provided that it can be carried out without ambiguity regarding the possible loss or exchange of the label, protein labelling using a fluorophore or a radioactive isotope built into the structure is a useful way of being able to identify the presence of a particular protein fairly unambiguously. On the other hand, using such label-based detection techniques as total internal reflexion fluorescence microscopy (TIRF), it is not possible to determine the absolute number of adsorbed molecules with precision. A further possible approach is to use infrared spectroscopy, preferably using a waveguide configuration as in OWLS. This can allow fairly reliable estimates of the presence of certain characteristic chemical groups such as carbonyl, but on the other hand most of the detectable groups are ubiquitous among proteins, and hence cannot be used to identify a particular protein; hence this method only be applied for identification purposes if some very exotic group is present in sufficient quantity.

7.4.2

Electrical surface properties

The ionization of surface chemical groups, such as surface hydroxyls, can engender an electrostatic charge at the surface of a solid. Alternatively or additionally, the charge can arise through the adsorption of ions from solution. The charge in the plane of the surface is compensated by the diffuse ionic atmosphere of counter ions. When the solution and the solid move relative to each other, the surface charge becomes separated from its counter ions, and the system defined by the solid object and its immediate environment loses its electrical neutrality. This phenomenon gives rise to the four electrokinetic phenomena. In electrophoresis, the solid moves (relative to the laboratory coordinates) and liquid remains stationery. In electro¨osmosis, the solid is fixed, and the liquid is moving. Liquid forced through fine capillaries engenders the so-called streaming potential between the two ends of the capillaries, and a particle accelerated through a liquid medium acquires an electric charge, called the sedimentation potential or Dorn potential. When a solid moves relative to its liquid medium, the plane separating the motions does not coincide with the boundary between the solid and liquid phases. Hence, the potential at the so-called slipping plane is invariably less than the electrostatic potential at the phase boundary. The potential at the slipping plane has a special name, the ζ-potential.

7.4. CHEMICAL SURFACE ANALYSIS

7.4.3

135

Protein adsorption and desorption

The non-ergodicity of most systems involving proteins places great emphasis on obtaining high-quality kinetic data in order to adequately experimentally characterize protein adsorption and desorption processes. Historically, a great many methods have been brought to bear; nowadays, heavy prominence is given to high resolution in situ methods, with which the absolute number of adsorbed particles can be determined. The two main methods used for this purpose are mechanical (based on the change of the resonant frequency of a vibrating cantilever or crystal due to the adsorption of matter on the surface of the cantilever or crystal) and optical (based on the change in reflectance of a thin interface between the adsorbent and contacting liquid due to the accumulation of adsorbate). In biomaterials research, one of the most important aspects is the ability to recreate the relevant surface in the required form for the measurement. In principle the vibrating cantilever can be made from, or coated with, any material. The vibrating piezoelectric crystal must be coated with conducting electrodes to which an alternating voltage of appropriate frequency is applied; the electrode could in turn be coated with the required material. Direct scanning angle reflectometry (SAR) or ellipsometry can be applied to any surface. The accuracy of ellipsometry is potentially considerably greater than that of SAR. Optical waveguide lightmode spectroscopy (OWLS), using grating couplers or interferometry, can be applied whenever the biomaterial can be fabricated in the form of a waveguide, or applied as a coating to an optical waveguide. The mechanical methods work very well in vacuum or gas, but in a liquid, the coupling of the vibrating solid with the fluid makes it difficult to interpret the results quantitatively. when trying to determine the adsorption of proteins from an aqueous solution. The optical methods will not work if the electronic polarizability of the absorbing particle is the same as that of the medium in which it is dissolved or suspended. There is a large family of optical techniques using evanescent waves generated at the the interface between two dielectrics, e.g. solid/liquid or solid/gas. In order of historical development, they are reflectometry, ellipsometry, surface plasmon resonance (SPR), and optical waveguide lightmode spectroscopy (OWLS) (see Figure 7.5). Each of them has their advantages and disadvantages. Reflectometry and ellipsometry are the most versatile in terms of the surfaces with which they can be used, there being essentially no restriction on opacity. Reflectometry has the advantage that the interpretation of the results involves a straightforward application of Fresnel’s equations, but it is slow and the achievable precision is low. Ellipsometry can be used in both reflexion and transmission mode, and has higher precision, but

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there is no closed-form solution of the governing equations, and therefore in order to extract surface parameters of interest a tedious fitting procedure has to be applied, with no guarantee of a unique result at the end. Furthermore, it is notoriously poor for the determination of the thickness and refractive index of surface layers, which are often of great interest in biocompatibility studies; its main use is to determine the total amount of adsorbed material. Surface plasmon resonance spectroscopy is restricted to investigating noble metal surfaces, for which it is more sensitive than ellipsometry. Nevertheless, since the surface plasmons only exist in the p-polarization, the amount of information that can be extracted from the measurements is limited. opt Figure 7.5: Characteristics, and advantages and disadvantages, of the various optical techniques available for thin surface film investigation.

The most recent technique is optical waveguide lightmode spectroscopy (OWLS), in which the surface under investigation is made part of an optical waveguide, and the actual measured parameters are the propagation constants of the lightmodes guided therein. The technique is that least an order of magnitude more sensitive than SPR, and both s- and p-polarizations can be measured. There is a straightforward relation between the measured parameters and the opto-geometrical parameters of the surface, which can therefore be determined with high precision and reliability. The technique is extremely valuable for studying the adsorption of proteins and nanoparticles at surfaces, for which it has been extensively used. Another strength of the technique is its versatility, since the optical waveguides can be made out of any high refractive index transparent dielectric material, or coated with a thin (∼ 10 nm) film of the material whose biocompatibility is under investigation. The investigation of opaque materials such as steel may be problematical: in this case a nanoparticle monolayer of the metal, or its oxide, can be assembled on the waveguide surface in order to approximate the opaque bulk. Various approaches have been adopted for measuring the lightmode spectrum: the two main ones are via grating coupling, and interferometry in one form or another. The use of OWLS for measuring the behaviour of living cells at biomedical surfaces has also been reported. In all these applications, it is the very high sensitivity, excellent time resolution and multiplicity of measurable outputs interpretable as real physico-chemical parameters of the system that give it an unchallenged advantage for the quantitative investigation of biocompatibility. Of particular value in far-reaching and fundamental

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studies is the possibility of reliably extracting the thermodynamic and kinetic parameters characterizing protein adsorption, as discussed Chapter 1.

7.5

Cytometry

We have already noted, as a tenet of the field, that the adsorption of proteins to candidate biomedical surfaces is fundamentally the most important aspect. The study of protein adsorption is indeed sufficient for establishing the biocompatibility of devices intended to be placed in the bloodstream. Furthermore, we have argued that since the interaction of living cells with biomedical surface is always mediated by proteins, the study of protein adsorption goes a long way towards understanding the response of living cells to such surfaces. Nevertheless, for implants designed to be in medium or long-term contact with living tissue, it is important to study the interaction of the cell as a unit with the artificial surface. Later on, one can explore the nature of the cellsurface contact in molecular detail. The basic phenomenology of cell-surface interaction comprises (i) adhesion (analogous to the adsorption of a protein molecule, cf. Figure 1.1); (ii) spreading (analogous to the denaturation of an adsorbed protein molecule), which formally means a change shape from (roughly) spherical to (roughly) that of a segment, implying a huge increase in cell-surface contact area (but note that the actual points of contact between the cell and the surface may be confined to quite small focal adhesions, cf. Chapter ??); growth (which has no direct analogy in protein adsorption; it could however be thought of as analogous to the accretion of dissolved proteins two a (denatured) adsorbed protein (cf. prion filament formation); and (iv) differentiation, which also has no analogy in protein adsorption. “Analogy” here means that effectively the cell is considered to be a giant macromolecule; the fact that the cell and especially the cell surface is partly made up of proteins is of no relevance. In the early days of the quantitative investigation of cell-surface interactions, the optical microscope was the main tool available, and has the very important advantage that in situ observation is possible essentially without modification of the experimental system. The disadvantage is that it is a fairly cumbersome tool, and can only be used “manually”, i.e. automation of observation and quantification is scarcely practicable. Nevertheless, excellent work was carried out.10 A more recent development involving optical microscopy has been to use total internal reflexion at the interface between a 10

A good example is the paper of A.C. Taylor, Attachment and spreading of cells in culture, Experimental Cell Research, Suppl. 8 (1961) 154–173.

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glass substrate and aqueous medium to excite the fluorescence of molecules specially introduced into the cell. By appropriate choice of the parameters of the experimental arrangement, the evanescent field generated by the total internal reflexion decays within a distance ∼ 100 nm.11 This has been particularly useful for mapping the focal adhesions. As with all fluorescencebased techniques, it has the disadvantage of requiring the introduction of non-physiological molecules (the fluorophores) into the system, and adequate verification of equivalence of behaviour of fluorophore-free and fluorescently labelled cells has not yet been carried out. By making the cell substrate electrically conducting, the electrical impedance of the cell-substrate contact can be measured. In the simplest realization, either the entire substrate is made conducting or it is decorated with microelectrodes, and the counterelectrode is present elsewhere in the medium, which is usually sufficiently well-conducting so that the entire current between the electrode and the counterelectrode passes through it. The impedance therefore gives a direct measure of the (insulating) cell-substrate contact area. Nevertheless, it is not usually possible to accurately determine that area from the impedance measurement alone. The system is electrically extremely complicated, especially when it is borne in mind that the cell-substrate contact is not a simple blocked area (Figure 7.6). A start has been made on determining the influence of frequency, amplitude and bias on the impedance, but given that the system is living and hence potentially constantly evolving, progress towards making the technique a truly quantitative one has been very slow. Furthermore, since in order to obtain reliable results the system needs to be calibrated by microscopic observation for practically every measurement series, there appears at present to be little advantage in carrying out electrical measurement to determine the biocompatibility; it is rather a tool for exploring the responses of cells to environmental stimuli, including electrical ones. These responses could of course include the production of proteinacious microexudates. The same comments apply to measurements made by printing interdigitated electrodes onto an insulating substrate, and measuring the impedance between them. It will be influenced in subtle ways depending not only on the overall cell-substrate contact area, but also on the details of the cell-substrate interface. The quartz crystal microbalance has also been used in to look at cell 11

D. Gingell et al., General electromagnetic theory of total internal reflexion fluorescence: the quantitative basis for mapping cell-substratum topography, J. Cell Sci. 87 (1987) 677–693; J.S. Burmeister et al., Quantitative analysis of variable-angle total internal reflexion fluorescence microscopy of cell-substrate contacts, J. Microsc. 173 (1994) 39–51.

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Figure 7.6: Sketch of the types of contact between a cell and a solid planar substrate. a, a cell cultured in bulk medium that has just landed on the surface, and still retains its spherical shape; b, an idealized spread cell of the same volume as its spherical precursor, and making closely parallel contact but the substrate; c, a more realistic representation of a spread cell, showing the somewhat irregular shape of the cytoplasm, and the presence of filopodia some tens of nanometres long between the main body of the cell and the substrate. The DNA-rich nucleus of the cell is represented by the hatched zone.

adhesion.12 With the device, the main problem is interpreting the results in terms of physically meaningful variables such as cell-substrate contact area. The evanescent field exploited in the total internal reflexion fluorescence microscopy technique described above has also been exploited in the application of optical waveguide lightmode spectroscopy (OWLS) to cytometry. By making the substrate an optical waveguide, large numbers of total internal reflexions are effectively taking place within the contact of a single cell. The quantitative interpretation of the results rely on the fact that the refractive index of the cell material is different from that of the surrounding medium. Hence the presence of the cell, and any changes of its shape, perturbs the evanescent field, which is constantly monitored via the propagation constants of the guided lightmodes in the waveguide. Since the evanescent field decays exponentially into the medium, one can very conveniently define an effective volume of the cell, which is simply the Laplace transform of its cross section parallel to the plane of the substrate. Provided the volume of the cell is known, measurement of even the propagation constant on the single mode is sufficient to determine the cell-substrate contact area with nanometric precision.13 An interesting extension of this idea makes use of the fact that the DNArich nucleus has typically a significantly different refractive index from that of the rest of the cytoplasm. Therefore, all other parameters being equal, any movement of the nucleus within the cytoplasm will also perturb the evanescent field. The problem in practice is to ensure that a sufficient number of lightmodes are measured in order to unambiguously separate effects of 12

C. Fredriksson et al., The piezoelectric quartz crystal and dissipation sensor: a means of studying cell adhesion, Langmuir 14 (1998) 248–251. 13 J.J. Ramsden et al., Cytometry.

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shape change and nuclear movement. The way that OWLS is typically exploited in cytometry is to firstly measure the rate of attachment of cultured, hence spherical, cells suspended in medium to a planar substrate. The key variable in the experiment is the wall shear rate of the flowing medium; it should be chosen such that the cells that actually remain on the surface have made an adhesive contact with it. The subsequent spreading takes place over much longer timescales. Detachment can be studied by simply gradually increasing the wall shear rate. In this case, OWLS simply serves as a counter of cells present at the surface of the substrate. The force required to attach a cell was one of the earliest way is used to characterize cell-substrate adhesion.14 Microtechnology has been used to construct a miniature shearing device in which a microcantilever is used to apply a lateral load to the cell.15 Cell viability can be tested using the usual methods of testing for the permeability of the cell membrane to a dye. Microphysiometry may be appropriate for more detailed studies of cell health on surfaces. Essentially, this means the measurement of extracellular acidification using some kind of miniature pH sensor. The most complete records of cell activity are obtained through studying repertoires of expressed proteins and metabolites.

7.5.1

Microarrays

Active and passive arrays. In the former, each element is a miniature, but functionally complete, biosensor. The motivation for the array is to achieve high throughput at low cost. In the latter, each element has a severely reduced function, and typically can only detect one particular thing. Reading off the result is carried out in a separate, off-line process. Identifying and quantifying the amounts of all the messenger RNA (mRNA) in a cell is done using microarrays (‘gene chips’). The principle of a microarray is to coat a flat surface with spots of DNA complementary to the expressed mRNA, which is then captured because of the complementary base pairing (hybridization) between DNA and RNA (A–U, C–G, G–C, T–A) and identified. The relationship of a microarray to a classical affinity assay resembles that of a massively parallel processor to classical linear processor architecture, in which instructions are executed sequentially. 14

A shearing machine is described by L. Weiss, The measurement of cell adhesion, Experimental Cell Research, Suppl. 8 (1961) 141– 153. 15 A. Yamamoto et al., A new technique for direct measurement of the shear force necessary to detach a cell from a material, Biomaterials 19 (1998) 871–879.

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Microarrays consist of a two dimensional array, typically a few square millimeters in overall area, of more or less contiguous patches, the area of each patch being a few square micrometres (or less), and each patch on the array having a different chemical composition. Typical microarrays are assembled from one type of substance (e.g. nucleic acid oligomers). In use, the array is flooded with the sample whose composition one is trying to identify. After some time has elapsed the array is scanned to determine which patches have captured something from the sample. Each patch is addressible, in the sense that the composition of each individual patch is known or traceable. Hence a photomicrograph of the array after exposure to the analyte should allow one to determine which substances have been captured from the sample. Detailed protocol for a microarray assay: Array preparation The chip must be designed on the basis of what one is looking for. Each gene of interest should be represented by at least one, or preferably more, unique subsequences. Once the set of sequences has been selected, there are two main approaches to transfer them to the chip: 1. Assemble successive monomers using microfabrication technology. For example photoactivatable nucleic acid monomers are prepared. Masks, or a laser scanner, activate those patches selected to receive, say, G. After exposure to light, the array is then flooded with G. Then the array is exposed to a different pattern, and again flooded (with a different base), and so on. This technology is practicable up to about 20 cycles, and is highly appropriate wherever linear heterooligomers sharing a common chemistry are required [?]. 2. For all other cases, minute amounts of the receptor substances can be directly deposited on the array, e.g. using laboratory microrobotics technology. This is suitable for large macromolecules, such as proteins, or sets of molecules substances not sharing a common chemstry, or longer oligopeptides. Solutions of the different substances are applied using a type of inkjet technology. This process is sometimes called ‘dip-pen lithography’. In both cases, each patch can be uniquely identified by its Cartesian array coordinate. Sample preparation The raw material is processed to release or uncomplex the analytes of interest, and possibly partially purified. The

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Array exposure . The array is flooded with the sample and allowed to reach equilibrium. Then, all unbound sample is washed away. Array reading The array is scanned to determine which patches have captured molecules from the sample. If the sample molecules were prelabelled with a fluorophore, then fluorescent patches indicate binding, with the intensity of fluorescence giving some indication of the amount of material bound, which in turn should be proportional to the amount of mRNA expressed in the original sample. The binding of unlabelled samples may be detected by certain compounds which fluoresce when intercalated into double stranded DNA, i.e. the spots hybridized by capture from the sample. Image processing The main task is to normalize the fluorescent (or other) intensities. It is important when comparing the transcriptomes from two samples (e.g. taken from the same tissue subject to two different growth conditions). A straightforward procedure is to assume that the total amount of expressed mRNA is the same in both cases (which may not be warranted, of course) and to divide the intensity of each individual spot by the sum of all intensities. If the transcriptomes have been labelled with different fluorophores and exposed simultaneously to the same chip, then normalization corrects for differences in fluorescence quantum yield, etc. Analysis The results are analysed using pattern recognition techniques.

7.5.2

The polymerase chain reaction

If the amount of DNA to be investigated is very small, it can be multiply copied (‘amplified’) by the polymerase chain reaction (PCR) before further analysis. The following steps are involved: 1. Denature (separate) the two strands at 95 ◦ C; 2. Lower the temperature to 60 ◦ C and add primer, i.e. short synthetic chains of DNA which bind at the beginning (the s-called 30 end) of the sequence to be amplified;

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3. Add DNA polymerase (usually extracted from the thermophilic microbe Thermus aquaticus and hence called Taq polymerase) and deoxyribonucleose triphosphates (dNTPs, i.e. an adequate supply of monomers); the polymerase synthesizes the complementary strand starting from the primer; 4. Stop DNA synthesis (e.g. by adding an auxiliary primer complementary to the end of the section of the template to be copied); go to step 1. The concentration of single strands doubles on each cycle up to about twenty repetitions, after which it declines: errors accumulate (in a living organism, sophisticated proof-reading and repair mechanisms ensure that the process can continue essentially indefinitely). Miniature bioMEMS (lab-on-a-chip) devices are now available for PCR.

7.5.3

DNA sequencing

For determining the ‘text’ of genes. The classical technique is that devised by Sanger. One starts with many single-stranded copies of the unknown sequence, to which a known short marker sequence has been joined at one end. An oligonucleotide primer complementary to the marker is added, together with DNA polymerase and nucleotides. A small proportion of the nucleotides are fluorescently labelled dideoxynucleotides lacking the hydroxyl group necessary for chain extension. Hybridization of the primer to the marker initiates DNA polymerization templated by the unknown sequence. Whenever one of the dideoxynucleotides is incorporated, extension of that chain is terminated. After the system has been allowed to run for a time, such that all possible lengths have been synthesized, the DNA is separated into single strands and separated electrophoretically on a gel. The electropherogram shows successive peaks differing in size by one nucleotide. Since the dideoxynucleotides are labelled with a different fluorophore for each base, the successive nucleotides in the unknown sequence can be read off by observing the fluorescence of the consecutive peaks. A useful approach for very long unknown sequences (such as whole genomes) is to randomly fragment the entire genome (e.g. using ultrasound). The fragments, approximately 2 megabases long and sufficient to cover the genome five to tenfold, are cloned into a plasmid vector,16 inserted into a bacterial genome and multiplied. The extracted and purified DNA fragments are then sequenced as above. The presence of overlaps allows the 16

In this context, ‘vector’ is used in the sense of vehicle.

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original sequence to be reconstructed. This method is usually called shotgun sequencing.

7.5.4

Two dimensional gel electrophoresis (2DGE)

This technique is used to separate proteins. Physico-chemical attributes of proteins are: 1. Molecular weight Mr ; 2. Net electrostatic charge Z (as a function of pH—the pH at which Z = 0 (isoelectric point) is important as a characteristic parameter). Both can be calculated from the amino acid sequence (assuming no postranslational modifications), provided Mr and Z of the individual amino acids are known. Mr is easy; to calculate Z one has to make the quite reliable assumption that all the ionizable residues are on the protein surface. The calculation is not quite as simple as adding up all the surface charges, since they mutually affect each other (cf. the surface of a silicate mineral: not every hydroxyl group is ionized, even at extremely low pH).17 2DGE was independently developed by Klose and O’Farrell in 1975. It depends on the fact that separation by isoelectric point (i.e.p. or pI, also referred to as point of zero charge, p.z.c.) is insufficient to separate the very large number of proteins found together in a typical living cell, many of whose i.e.p.s are clustered together. Equally, there are many proteins with similar molecular masses. By applying the two techniques sequentially, however, they can be separated, especially if large (30 × 40 cm) gels are used. Proteins in the crude cell extract are dispersed in an aqueous medium containing the anionic detergent sodium dodecyl sulphate (SDS) which breaks all noncovalent bonds, i.e. subunits are dissociated; the first separation takes place according to isoelectric point by electrophoresis on a gel along which a pH gradient has been established; the partly separated proteins are then transferred to a second, polyacrylamide, gel within which separation is effected according to molecular weight. If the cells have been pulse radiolabelled prior to making the extract, then the final gel can be scanned autoradiographically and the density of each spot is proportional to the net rate of protein synthesis. Alternatively the proteins can simply be stained and the gel scanned with a densitometer; 17

Linderstrøm-Lang worked out a method of taking these correlations into account; his formula works practically as well as more sophisticated approaches (including explicit numerical simulation by Brownian dynamics) and is much simpler and more convenient to calculate.

7.6. QUANTIFYING BIOCOMPATIBILITY

145

the spot density is then proportional to protein abundance. The molecular mass and isoelectric point ranges are limited by practical considerations. Typical ranges are 15 000 < Mr < 90 000 and 3 < pI < 8.

7.5.5

Proteolysis and chromatography

A technique for separating proteins: enzymatic proteolysis cuts them into shorter peptides followed by column chromatography. Trypsin is usually used as the proteolytic enzyme since it cuts at well-defined positions (lysines). The principle of chromatography is to functionalize a stationary solid phase (granules of silica, for example) packed in a column and pass the sample (suspended or dissolved in the liquid mobile phase) through it. The functionalization is such that the proteins of interest are bound to the granules, and everything else passes through. A change in the liquid phase composition then releases the bound proteins. Better separations can be achieved by “multidimensional” chromatography, in which a cation exchange column (for example) is followed by a reverse phase column. The number of dimensions can be increased further.18 Usually the technique is used to prepurify a sample, but in principle, using differential elution (i.e. many proteins of interest are bound, and then released sequentially by slowly increasing pH, or the polarity of the liquid) can be also used to separate proteins prior to identifying them individually.

7.6

Quantifying biocompatibility

It should by now be clear that biocompatibility is not a simple universal variable such as, for example, mass or electrostatic charge. Furthermore, in biomedical applications the designer of a device to be implanted usually only wishes to know whether the material he or she proposes to use is biocompatible or not. The answer to be given must moreover depend on the context (i.e. the environment) of the application. The typical sorts of tests that are carried out our to measure the effective removal of active platelets from the blood, to measure the deposition of platelets on the surface, and to measure the activation of platelets. Similarly, the adsorption of selected proteins (fibrinogen is a favourite, but some of the clotting factors are also investigated) has also been investigated. Finally, haemolysis is also considered as a biocompatibility marker. 18

2-dimensional gel electrophoresis (§7.5.4) is essentially an example of this.

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7.6.1

Effects of roughness on response

Systematic data on the influence of roughness—at all relevant length scales— on the responses evoked in biological milieux is lacking. Since interfacial energy is strongly curvature-dependent, the distribution of curvatures, another roughness parameter, is likely to be very important, but again there is a dearth of data. The rule of thumb seems to be that implants destined for the bloodstream should be as smooth as possible, in order to discourage protein adsorption, and implants incorporated into tissue should be rather rough, to facilitate cell anchorage. Now that it is technically possible to machine metal and other surfaces to a roughness (or rather smoothness) of the order of 1 nm, it is very important to know whether the cost of such ultra precision machining is justified through superior performance.

7.7

Further reading

F.P. Bowden and D. Tabor, Friction and Lubrication of Solids. London: Oxford University Press (1954). de Chiffre, L., Lonardo, P., Trumphold, H., Lucca, D.A., Goch, G., Brown, C.A., Raja, J. and Hansen, H.N., 2000, Quantitative characterisation of surface texture, Annals CIRP, 49/2, 635652. Goch, G., Prekel, H., Patzelt, S., Strbel, G., Lucca, D.A., Stock, H.-R. and Mehner, A., 2004, Non-destructive and non-contact determination of layer thickness and thermal properties of PVD and sol-gel layers by photothermal methods, Annals CIRP, 53/1, 471474. Hillman, P., 1991, The Case for New Paradigms in Cell Biology and in Neurobiology. Lewiston: Edwin Mellen Press. Hocken, R.J., Chakraborty, N. and Brown, C., 2005, Optical metrology of surfaces, Annals CIRP, 54/2, 705719. Lehmann, P., Goch, G., 2000, Comparison of conventional light scattering and speckle techniques concerning an in-process characterization of engineered surfaces, Annals CIRP, 49/1, 419422. Lehmann, P., Patzelt, S., Schne, A., 1997, Surface roughness measurement by means of polychromatic speckle elongation, Appl. Optics, 36, 21882197. P. Lundahl, A. Lundqvist and E. Greijer, eds, Quantitative Analysis of Biospecific Interactions, Amsterdam: Harwood, 1998. Mazusawa, T., 2000, State of the art of micromachining, Annals CIRP, 49/2, 473488 Moore, J.H. and Spencer, N.D. (eds) (2001). Encyclopaedia of Chemical Physics and Physical Chemistry. Philadelphia: IOP.

7.7. FURTHER READING

147

Ramsden, J.J. Kinetics of protein adsorption. In: M. Malmsten (ed.), Biopolymers at Interfaces. New York: Dekker (1998), Ch. 10 (pp. 321–361). Ramsden, J.J. Adsorption Kinetics of Proteins. In: Encyclopaedia of Surface and Colloid Science, A. Hubbard (ed). New York: Dekker (2002) pp. 240–261. Ramsden, J.J. Protein adsorption kinetics. In: M. Malmsten (ed.), Biopolymers at Interfaces, 2nd Edn. New York: Dekker (2003), Ch. 8 (pp. 199–220). Ramsden, J.J., 1994, Experimental methods for investigating protein adsorption kinetics at surfaces, Q. Rev. Biophys., 27, 41-105. Ramsden, J.J., Li, S.-Y., Heinzle, E. and Prenosil, J.E., 1995, An optical method for the measurement of number and shape of attached cells in real time, Cytometry 19, 97102. Ramsden, J.J., 2006, OWLSa versatile technique for drug discovery, Frontiers Drug Design Discovery 2, 211223. Rodahl, M., Hk, F., Krozer, A., Brzezinski, P. and Kasemo, B., 1995, Quartz crystal microbalance setup for frequency and Q-factor measurements gaseous and liquid environments, Rev. Scientific Instruments, 66, 39243930. Rosencwaig, A. and Gersho, A. Theory of the photoacoustic effect with solids, 1976, J. Appl. Phys., 47, 6469. Sirohi, R.S. (2002). Speckle interferometry. Contemp. Phys. 43, 161– 180. Smith, G.T., 2002, Industrial Metrology: Surfaces and Roundness, Heidelberg: Springer. Teague, E.C., Vorburger, T.V. and Maystre, D., 1981, Light scattered from manufactured surfaces, Annals CIRP, 30/2, 563570. Vickerman J.C., ed., 1997, Surface Analysisthe Principal Techniques, New York: Wiley. Whitehouse, D.J., 1994, Handbook of Surface Metrology, Bristol and Philadelphia: IoP.

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Chapter 8 Prospects for smart and nano materials One of the most promising areas of research into biomaterials surfaces is ways in which such surfaces can be made “smart”, i.e. responsive to their environment. The notion of information flowing from the surface to cell has already been elaborated in Chapter 4, and will now be extended to cover information flowing from cells or macromolecular assemblies to the surface, which adapts appropriately.

8.1

The nature of responsiveness

Smart coatings essentially mean responsive coatings, i.e. they can be described as machines with input, and their output depends on their present state and their input. There are already several sensing or indicating coatings with nanoadditives available based on this principle. Examples are the oxygen indicating materials referred to above, and coatings incorporating piezoelectric particles giving off electrical pulses when deformed. Figure 8.1 shows the general nature of a responsive system. cyber Figure 8.1: Diagram showing information flows in a general responsive system.

Many of the smart coatings discussed in the literature rely on ionizable groups in the polymer becoming charged or uncharged, whereupon they will 149

150CHAPTER 8. PROSPECTS FOR SMART AND NANO MATERIALS collapse or form a fairly stiff comb, resectively (Fig. 8.2). These fairly dramatic structural changes are evoked by changes in the acidity or alkalinity of the solution. Similar principles are used to design porous materials, the orifices of which depend on the alignment of the pH-sensitive molecules. As yet, however, there has been litle progress in developing this kind of functionality in solvent-free systems. PEL Figure 8.2: Sketch of polyelectrolyte chains tethered by their ends to a substrate (hatched). (a) collapsed state; (b) extended state induced by adding a strong acid HA, which ionizes and protonates the amine groups, causing them to become electrostatically charged.

8.2

Bioactive coatings

The concept here is in effect to transform biomedical implants into pharmaceuticals, such that they actively intervene in the physiology of the host organism. The kinds of activities that are of interest are the prevention of blood coagulation, the encouragement of endothelial cell attachment, the prevention of excessive smooth muscle cell proliferation, etc. These activities go beyond what might be considered to be the basic functions of implants, e.g. joint replacement, and keeping a blood vessel open, but the effectiveness of a purely passive implant is likely to be rapidly vitiated by undesirable physiological responses. In other words, there is no such thing as a truly passive implant; its mere presence will trigger extraordinary, and usually undesirable, physiological activity around it, and the purpose of bioactive coatings is to steer that extraordinary activity back to ordinary activity, and in some cases to promote extraordinary but beneficial activity. The antimicrobial silver ion-releasing coatings mentioned earlier are a successful realization of the concept, at least within a limited frame of reference (adequate studies on the toxicity of the silver to the host organism do not yet seem to have been carried out); silver is also used to dope polymeric materials used for catheters. The two main ways in which bioactivity is achievable are: 1. coating with fragments of biomacromolecules (polypeptides or polysaccharides); and

8.2. BIOACTIVE COATINGS

151

2. incorporating a reservoir of small molecules (drugs, hormones, growth factors etc.) within the implant, most typically as a thick porous coating, from which the small molecules are slowly released during the lifetime of the implant. (1) is very useful for promoting the adhesion of cells to an implant. A synthetic material such as polyethylene will, in an uncoated state, first of all be coated by biopolymers adsorbing from the bloodstream or secreted by cells in its vicinity. At present, we lack sufficient knowledge to effectively control this process; therefore the deliberate coating of the implant with biomacromolecular fragments known to elicit a specific response is a very practical way for achieving some degree of control over the organism’s response to the presence of the implant.

8.2.1

Cell adhesion-promoting coatings

Some of the most extensively researched adhesion-promoting coatings are based on the tripeptide motif arginine-glycine-aspartic acid (RGD), which is a fragment of the ubiquitous adhesion protein fibronectin. RGD is a receptor for a particular integrin present on the surface of endothelial cells. It is much more effective to coat the implant with the RGD fragment than with the whole protein, firstly because the receptor density is much higher (and the possible problem of the denaturation of the mobilized entire protein is avoided), and secondly because the coating is then highly specific for endothelial cells: fibronectin contains many other fragments acting as receptors for other kind of cells.

8.2.2

Call adhesion-inhibiting coatings

Coatings that can slowly release or generate in situ bioactive substances are very promising as biocompatibilizing agents. For example, the simple molecule nitric oxide, NO, is a well-known inhibitor of platelet adhesion and activation. It is therefore a candidate for incorporation into the surface of a stent, for example. Endothelial cells are estimated to release about a million molecules of NO per second [Frost], which would therefore be the quantitative target for the drug-releasing material. Possible approaches include blending a small molecule NO-donor into a polymer, covalently binding the NO-donor to polymer side chains, and covalently binding the donor to the polymer backbone. Bioactivity also includes the ultimate in implant assimilation—the implant that is totally degraded and metabolized or excreted by its host after

152CHAPTER 8. PROSPECTS FOR SMART AND NANO MATERIALS it has fulfilled its function. The most promising approaches to this goal are materials that do not simply decompose in a predetermined fashion, but which react according to signals (e.g. secreted enzymes) emitted by the cells surrounding them.

8.2.3

Drug delivery surfaces

Drug delivery materials are typically nano- or microporous. Some of the general properties of this type of material can best be considered making use of percolation theory and anomalous diffusion. The interaction of drug delivery particles that circulate within the body follows directly from the principles in earlier chapters. Advanced catheter coatings involve the use of antibiotics such as Triclosan to impregnate the silicone from which the catheter is made in order to discourage bacterial colonization (especially by P. mirabilis) of the catheter. Porosity Nanoporous nanoparticles are extremely useful vehicles for the controlled release of drugs, perfumes and other volatile substances, either to the interior or the exterior of a package. Conversely, nanopores will also enhance the scavenging function of nanoadditives by further increasing the surface area. The control of the porosity to provide a desired dynamic release profile is in its infancy. Considerations of diffusion on non-Euclidean lattices will apply, i.e. the so-called anomalous diffusion. A further consideration is the affinity of the volatile substance to the material of the nanoparticles (or the substance with which the pores are coated), and how it can be modulated.

8.2.4

Biocidal surfaces

The main type of biologically active coating in current use is the antimicrobial coating. These coatings must merely secrete a toxin for the microbe of interest, often bacteria. Silver ions are toxic to most cells, including many bacteria, and matrices containing dispersed metallic silver particles slowly release silver ions. Here nanoparticles offer a clear cost advantage, since the total surface area determines the rate of release of the ions; hence for a given rate of release, less costly silver (or other antimicrobial agent) is required. The nanoparticles will of course be exhausted sooner than microparticles, but will in nearly every case outlast the design life of the object being protected. Nanoporous particles loaded with an organic toxin may be included in the coating formulation.

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153

In another type of coating, cellular adhesion is simply inhibited, and may thereby prevent cellular proliferation. The exact role of the nanoadditives is not so clear in this case, beyond allowing an unusual combination of material qualities. Antimicrobial coatings have a variety of functions, ranging from protection against disagreeable odours to killing pathogens in operating theatres. If illumination is feasible, photocatalytic particles are typically more effective than those operating mainly through thermal effects.

8.3

Carbon nanotubes in biomedical surfaces

Carbon nanotubes are one of the most exciting children of the nanorevolution. Carbon nanotubes are large molecules of pure carbon that resemble rolled up sheets of graphite. The tubes are approximately one nanometre in diameter. Despite, or perhaps because of, their minuscule size they display remarkable physical and chemical properties. They are almost 100 times stronger than steel by weight, and are potentially the best conductors of heat and electricity ever discovered. These key features can be used to create a variety of new and improved materials. Carbon nanotube-based composites are being widely researched to develop ultra lightweight and extremely strong materials for aerospace and other high performance applications. Since their discovery in 1991 in Japan, carbon nanotubes have promised the development of a wide range of applications and novel products, but commercial research and development into these applications has been constrained by the lack of a reliable manufacturing process. The availability of commercially produced carbon nanotubes of a consistent quality and at an accessible price is expected to unlock the potential for a wide range of industrial applications including ultra-strong materials (lightweight plastics that are 100 times stronger than steel), conductive composites and new electrical and electronic devices. At present, from the biomedical viewpoint this appears to be the technology is looking for application. As in so many other by materials, however, it may be supposed that exaptation of successful electronics etc. applications in the biomedical sphere will occur.

154CHAPTER 8. PROSPECTS FOR SMART AND NANO MATERIALS

8.4

A complexity sciences perspective for biomaterials design

The surface of a cell—and this applies as much to the tissue cells in contact with a prosthetic implant as to the blood lymphocytes transiently in contact with a stent—is constantly relaying information about its chemical, physical and mechanical environment back to its “command centre”—the protein expression machinery mainly localized in the cell nucleus. Within the limitations of the likely modification of any artificial surface by its biological environment, this information channel opens up possibility of designing smart materials that can direct the work of the cell in secreting desirable macromolecules. The design concept involves ensuring that some appropriate molecule, or molecular motif, embedded in the biomedical surface interacts with a receptor on the surface of the cell, triggering internal changes that eventually lead to the activation of transcription of a protein, or the inhibition of transcription. That protein, in turn, could well be secreted at the surface of the cell, providing a kind of feedback. Given the complexity of the inflammatory response, there are numerous points of possible intervention, both morphological and chemical, available to the biomaterials designer. In fact, this aspect of biomaterials design has only just begun to be seriously investigated; knowledge is still highly fragmentary at present, and it has not proven easy to even establish the main attributes of foreign particles present within the body responsible for inducing an inflammatory response.

8.5

Further reading

J.J. Ramsden, Bioinformatics: and Introduction, Dordrecht: Kluwer (2004).

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