Medical Device Link.

MATERIALS

Physicochemical processes

The first encounter an implanted device has within the body involves its surface. Whatever mechanical or bulk-related attributes an implant has, it is almost certainly its surface properties that trigger the body’s initial acute, response. This response then conditions the later programmed, complex cascade of biological responses, which determine the degree to which the implant becomes tolerated or rejected. The body’s ability to recognise the alien construct is exceptional, it having been honed through evolution to recognise the “non-self.” Whatever artificial masking or device materials and packaging strategies are used, the body can discriminate the imposed difference. Even approaches that attempt to mimic biology do not fully overcome the combination of cellular and humoral defence. In this context, there is no “biocompatible” material, but modern approaches do help to stave off biological attack to extend in vivo operational lifetime.

Through redoubled efforts at surface refinement and testing, there is an increasing understanding of the fundamental physicochemical processes that drive the body’s response to an implant material. It is also now possible to better gauge the quantitative, relative importance of different surface modifiers. Much of this understanding centres on new, model designer surfaces, systematic investigation in controlled biofluids and the new measurement tools arising from surface-probe techniques such as atomic force microscopy, which are not constrained to vacuum conditions as were their predecessors.

Water at surfaces

In the vicinity of the surface, there are extreme solute concentration gradients involving colloids, ions and organics, which, though difficult to access, are the focus for important interfacial events. The conventional consideration of the bulk biomatrix reveals little of this dramatic and changing microenvironment. The polarised distribution of solutes, especially gradients in ionic charge, create a backdrop to many of the observed, dramatic surface biological effects. However, above all, it is the organisation of water molecules in this environment that may provide the answers to materials development needs. There is evidence that transition from bulk to surface bound water affects the stability and conformational unravelling of proteins near the interface. More protein-tolerant surfaces probably have loosely bound intermediate water layers that reduce direct protein interactions with this damaging surface water layer. This layer has a powerful capacity to disrupt protein and cellular hydration shells and, therefore, to create local denaturation/damage.

Figure 1: 3D protein arriving at an artificial rigid surface versus natural flexible membrane.

Modified surface chemistry

The pre-eminent surface modification layer is polyethylene glycol (PEG). It is well known that end-attached PEG resists surface biofouling. Its properties are thought to be based on its near-neutral state (low surface charge), its inherent flexibility and its ability to confer microscale structural flexibility to a solid surface. It is, however, likely that the real basis for its effects lies more deeply in its loosely associated water molecules, which are able to interact less destructively with biological macrostructures. Flexibility and interfacial “fluidity” is a characteristic of many biological interfaces, not least cell membranes. It is hypothesised that these constructs may reduce surface distortion of encountered proteins (Figure 1) and dampen this major denaturing trigger for a future host rejection response. The problem for biomaterials and biosensors is that in contrast to self-renewing biological motifs and interfaces, artificial materials need to be structurally robust, and built-in fluidity is difficult to achieve.

Self-assembled monolayers provide an excellent, reproducible model for the study of surface processes defined by modified chemistries. Techniques now make it readily possible to change, for example, chain length, terminal motifs, bridging molecules and packing density.¹ It is now becoming apparent from this work that microscopic molecular domains are the primary drivers for the quality of protein-surface and even cell-surface interactions. These are more important than global, macroscale properties, for example, hydrophobicity/hydrophilicy of any surface; proteins themselves have contrasting surface motifs and selective functional groups for specific surface docking. Resultant protein unravelling at a surface through surface competing binding forces has been shown to be rapid compared with the later lateral surface movement or segregation into different biomaterial surface domains. Surface engineering allows control over the degree of this movement and can achieve masking or exposure of more selective cell binding domains (Figure 2), for example, allowing directed cell movement along lateral surface gradients. The more powerful lateral effects will operate by stimulating reactive cell adhesion/de-adhesion sequences, which can then be a further control route to cell movement and cell guidance in vivo (not so relevant for proteins). At the opposite extreme, to control binding, surface repulsion can be engineered to some degree. This has been reported for sulphonated polyethylene oxide to create a “negative cilia” effect that is able to resist protein and platelet adhesion through direct charge effects.²

Figure 2 : Sterically hundred cell binding motif at A (discs) leads to a shift in cell attachment to B.

Polyelectrolyte layers

Alternating negatively and positively charged polymeric monolayers of natural and artificial polymers are now readily deposited on surfaces. Hyaluronic acid and poly(L-lysine), for example, have been used to create highly controlled oligo layers. This oligo layer-by-layer (l-b-l) technique is the horizontal equivalent of the vertical self-assembly of monolayers. It can be used to control, for example, the level of exposed biomaterial surface charge, and to thereby modulate Coulombic surface binding interactions with cells and colloids and to manipulate surface profile, notably roughness. Their derivatisation can allow tethering of added surface-active molecules and the interlayer voids can provide a sophisticated multi-environment reservoir for time phase controlled drug release (Figure 3), perhaps of an antimicrobial. An l-b-l chitosan/hyaluronic film has even demonstrated feasibility as a protective barrier for a damaged artery, being able to mask drivers of surface coagulation.³ Plasmid deoxyribonucleic acid within a degradable l-b-l film has been tried in controlled delivery for genetic modification of cells on biomaterials. Applying this chemistry as a controlled layer in the x-y plane as well as in the z axis is beginning to show mechanistically comprehendible pictures of what biology really does. As with early biological science, studies are now able to move from the empirical and descriptive to the predictive.

Surface biomolecules

Figure 3: Polyelectrolyte layers encapsulating a therapeutic agent for controlled release. Each layer can have variable density and structural organisation, provided there is charge balance.

Heparin is the archetype surface bio-modifier, with a plethora of near-ideal characteristics. It is nontoxic, has multi-system effects, is biologically potent and its effects are dose dependent. It primarily amplifies the anticoagulant potency of antithrombin. It also modifies surface adsorption of proteins and has various distinct effects on proteolytic enzymes, fibrinogen and complement, and modulates cell attachment, for example, reducing white cell adhesion. The latter is of specific benefit in scaling down tissue inflammatory response to a biomaterial. It is known that other biological modifiers exert their potent effects through complex, multi-system influences; they were after all engineered by evolution. More research effort is being made to use biomolecules and in many cases a new rationalisation of protective mechanisms has been achieved. In this work, polysaccharides based on alginates have been engineered to reduce cell attachment, even to stainless steel. Also, immobilised phosphorylcholines as zwitterionic molecules have proved to mimic the outer cell membrane and have conferred resistance to protein adsorption, despite loss of cell membrane molecular mobility. This is probably achieved by attenuating the disruptive properties of surface water. Integrin binding peptides, for example, bearing arginine-glycine-aspartate motifs, determine cell adhesion in a receptor-specific way. However, their surface presentation affects potency (Figure 2) and effects can be selectively masked by embedding in other coatings. More subtly, these can also be modified by inclusion in linear peptide sequences versus, for example, multi-branched oligomers, or variously employed in synergistic combinations with other biological motifs.

Tissue-related interactions

In tissue, beyond the implantation damage, the implant triggers a broad platform of intense surface reactions, which are delivered by disrupted and permeabilised blood capillaries. The initial events then recruit cell build-up. Essentially, this is the acute inflammatory sequence. Beyond this, the classical wound healing response is well known, what is missing is the link with the early cell–protein adhesion process. It would be surprising if surface, topography and micromechanics did not play a part in orchestrating the later cell-surface interactions. Certainly, exposed surface charge (mainly zeta potential) modulates the strength of initial cell adhesion, cell spreading; also surface mosaics, however random, condition the preferred cell contact points. Cell-surface interactions now also need to be studied using metabolic probes. Cell respiration and metabolic oscillatory behaviour will be drastically altered by surface contact and there is a need to move beyond morphological and histological characterisation of tissue responses at implant surfaces to functional (metabolic and genomic) descriptors.

Figure 4 : Keratincyte cells grown on polypyrrole loaded with (a) dermatan (b) chloride and (c) gold surface control. (Figure courtesy of Dr Davidson Ateh.)

Emphasis on conducting polymers

Charge induction at a redox polymer through voltage application (DC or AC) enables its surface to be electrically manipulated and cycled, and thereby a manipulation of the superimposed biological environment may be achieved. This type of work is at an early stage in biomaterials and electrical phenomena are regarded by some as being unimportant in this area. Yet, charge field effects could alter cell contact, shape and migratory behaviour with a direct link, for example, to the oxidation state of a redox or conducting polymer. Tissue electrical stimulation using microelectrodes is known to alter local accumulation of proteins and thereby cell attachment. In the oxidised (charged) state, redox polymers can also entrap counter ions, which, if biomolecular, can confer altered bioactivity. Polypyrroles (Figure 4) and polythiophenes are the most studied conducting polymers, to date, and have allowed different structural organisation and chemistries to be generated. Further microscale structural refinement and electrochemical deposition of these materials could also engineer orientated structures and anisotropy, thus enabling better cell guidance for tissue regeneration where precise tissue architecture is important.

Likely future developments

Beyond surface engineering of macrostructures, micro- and nanotechnology will bring an ability to refine surface motifs (chemical and topographical) to unparalleled high tolerances. Biology is well able to operate at this scale, approaching the molecular; thus, the new structures will be valuable model systems for studying the interplay between surface cues. Clinical biomaterials design is likely to be applied in a manner increasingly typical of therapeutic drug agent development.⁴ One proviso is that however much understanding of short-term events at the interface maybe extended in this way, it does not necessarily allow long-term outcomes with implants to be influenced. What may be needed beyond understanding of the short-term is a contribution from smart materials research, particularly for responsive drug release materials to enable pharmacological overriding of later adverse host responses. The potency of drug-eluting stents, at least for the first few months after insertion, signals this convergence of biomaterials and pharmacological technologies. The place of interfacial science in this new “solid therapeutics” era may be to help control the outward transfer of bioactive, therapeutic solutes rather than the inward movement of destructive host-generated solutes.

Professor Pankaj Vadgama is Director, IRC in Biomedical Materials, Queen Mary, University of London, Mile End Road, London E1 4NS, UK, tel. +44 20 7882 5285,
e-mail: p.vadgama@qmul.ac.uk, www.irc-biomed-materials.qmw.ac.uk.