Biomaterials

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This section defines biomaterials and biocompatibility, classifies the materials used in medicine with their advantages, limitations and applications, and covers the principal methods of surface characterization.

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Definitions • A biomaterial is any material, natural or synthetic, intended to interface with biological systems in order to evaluate, treat, augment or replace any tissue, organ or function of the body (Williams' consensus definition).

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The defining feature is not the material itself but its intended contact with living tissue. • Biocompatibility is the ability of a material to perform with an appropriate host response in a specific application.

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Two consequences follow, and both are examined: it is application-specific, not an intrinsic property — a material acceptable in a hip stem may be unacceptable in a blood-contacting catheter — and an appropriate response is not necessarily no response; a bioactive material is meant to provoke a bonding response. • Related terms: bioinert (minimal interaction, a thin fibrous capsule), bioactive (forms a chemical bond with tissue — bioglass, hydroxyapatite), bioresorbable/biodegradable (dissolves and is replaced by tissue — PLA, tricalcium phosphate), biotolerant (accepted with a thicker capsule), haemocompatible (does not damage blood or provoke thrombosis), osteoconductive (serves as a scaffold for bone growth), osteoinductive (actively recruits and induces cells to form bone) and osseointegration (direct structural and functional connection between living bone and an implant surface). • Requirements of a biomaterial: non-toxic, non-carcinogenic, non-immunogenic and non-thrombogenic; adequate and appropriate mechanical properties (strength, stiffness matched to tissue, fatigue and wear resistance); corrosion and degradation resistance; sterilisable without degradation; manufacturable and reproducible at acceptable cost; and available in the required form. • Testing follows the ISO 10993 series for biological evaluation — cytotoxicity, sensitisation, irritation, systemic toxicity, genotoxicity, implantation and haemocompatibility — proceeding from in vitro cell tests through in vivo animal implantation to clinical trial.

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Host Response and Material Response Host response to the material Material response to the host Inflammation and the foreign body reaction Corrosion of metals; ion release Fibrous encapsulation Degradation and hydrolysis of polymers Thrombosis and complement activation Wear and generation of debris Infection and biofilm formation Fatigue and fracture Toxicity, allergy and carcinogenicity Absorption of water, lipids and proteins; leaching of additives Calcification of implanted devices Stress corrosion cracking and creep Classes of Materials Used in Medicine Class Advantages Limitations Typical applications Metals High strength, ductility, toughness, fatigue resistance, electrical conductivity Corrosion and ion release, very high modulus causing stress shielding, high density, image artefacts Joint prostheses, bone plates, screws, nails, dental implants, stents, pacemaker cases, instruments Ceramics / glasses Very hard, inert, excellent wear and compressive strength, good aesthetics, bioactivity possible Brittle, low fracture toughness and tensile strength, difficult to fabricate Femoral heads and acetabular liners, dental crowns, bone graft substitutes, coatings Polymers Easily formed into complex shapes, flexible, low density, resorbable versions possible, low cost Low strength and modulus, creep and wear, degradation, leachable additives, may absorb water Sutures, catheters, vascular grafts, IOLs, bone cement, UHMWPE bearings, drug delivery, tubing Composit es Properties tailored, stiffness matched to bone, strong and light Complex and costly to make, interface debonding, difficult quality control Dental restorative resins, bone plates, prosthetic limbs, fibre-reinforced scaffolds Natural / biological Excellent biological recognition, similar to native tissue, remodel in vivo Variable supply and properties, immunogenicity, disease transmission risk, need for crosslinking Collagen and gelatin scaffolds, bioprosthetic heart valves, allografts, chitosan, alginate, silk, hyaluronic acid Methods of Surface Characterization Because cells and proteins meet only the outermost few nanometres of a device, surface analysis is more relevant to biocompatibility than bulk analysis.

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Methods are chosen by what they reveal, how deep they sample and how much they damage the sample.

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Method Measures Sampling depth Contact angle / wettability (goniometry) Surface energy and hydrophilicity; a low contact angle means a hydrophilic, high-energy surface.

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Simple, cheap and very sensitive to contamination ~0.3-2 nm XPS (ESCA) — X-ray photoelectron spectroscopy Elemental composition and chemical bonding state (oxidation states) — the workhorse of biomaterials surface analysis; detects all elements except H and He ~1-10 nm Auger electron spectroscopy (AES) Elemental composition with high lateral resolution; good for small features ~1.5-3 nm SIMS (static/ToF-SIMS) Molecular and elemental fragments with very high sensitivity, including hydrogen and organic species; semi-quantitative ~1-2 nm ATR-FTIR Functional groups and molecular structure by infrared absorption ~1-5 μm SEM (with EDX/EDS) Surface topography and morphology at high magnification;

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EDX adds elemental analysis.

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Often needs a conductive coating ~5 nm - 5 μm AFM / STM Three-dimensional topography and roughness at nanometre resolution; can work in liquid and can map adhesion and stiffness ~0.5-5 nm Profilometry (stylus or optical) Roughness parameters (Ra, Rq) over larger areas ~1 nm vertical Ellipsometry / QCM / SPR Thin-film thickness and adsorbed protein mass in real time — the standard tools for studying protein adsorption Sub-nm Zeta potential (streaming potential) Surface charge in an electrolyte — important for protein and cell adhesion Interface • Practical points: most electron and ion techniques require ultra-high vacuum, so hydrated biological surfaces must be dried or frozen — the surface analysed may not be the surface the tissue met.

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Surfaces also reconstruct in response to their environment, a hydrophobic polymer turning its polar groups outward in water.

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Always analyse clean, representative, unhandled samples, and combine at least two complementary methods.

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This section covers the structure, chemistry, mechanical properties and applications of the metals used as biomaterials, corrosion and its control, and the natural materials — collagen with its physical and chemical modification, proteoglycans and glycosaminoglycans.

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Metallic Biomaterials:

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Structure and Requirements • Metals are used wherever high strength, toughness and fatigue resistance under load are needed.

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They are crystalline, with FCC, BCC or HCP lattices and metallic bonding that gives ductility, thermal and electrical conductivity.

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Their properties are controlled by alloying, grain size (the Hall-Petch relation — finer grains give higher strength), work hardening and heat treatment. • Requirements for an implant metal: excellent corrosion resistance in a warm, aerated, chloride-rich electrolyte at pH 7.4 (and as low as pH 3-4 in inflammation); non-toxic ions; adequate strength and fatigue life (a hip stem sees about 10⁶ cycles a year); acceptable modulus; good wear resistance; and manufacturability.

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Alloy Composition and structure Properties Uses 316L stainless steel Fe with 17-20 % Cr, 12-14 % Ni, 2-3 % Mo, austenitic (FCC);

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L = low carbon (< 0.03 %) to prevent sensitisation (chromium carbide precipitation at grain boundaries) E ≈ 200 GPa; yield 200-700 MPa depending on cold work; cheapest and easiest to fabricate; least corrosion resistant of the three — susceptible to pitting and crevice corrosion Temporary devices — bone plates, screws, nails, wires, some stents, instruments Cobalt-chrom ium alloys Co-Cr-Mo (cast, ASTM F75) and Co-Ni-Cr-Mo (wrought);

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Cr ~27-30 %, Mo ~5-7 % E ≈ 210-230 GPa; very high strength and the best wear resistance of the implant metals; excellent corrosion resistance; hard to machine;

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Co and Ni ion release is a concern Femoral heads and stems, knee components, dental partial dentures, stents Titanium and Ti-6Al-4V CP Ti (grades 1-4, HCP α) and Ti-6Al-4V (α+β alloy); spontaneously forms a TiO₂ passive film E ≈ 110 GPa — the closest of the three to bone; low density (4.5 g/cm³); best corrosion resistance and best osseointegration;

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MRI-compatible; but poor wear resistance (notch-sensitive, galls) and higher cost; concerns over V and Al have led to Ti-6Al-7Nb and Ti-Nb-Zr Dental implants, hip stems, plates and screws, pacemaker cases, spinal cages Nitinol (Ni-Ti) Near-equiatomic Ni-Ti with a reversible martensite-austenite transformation Shape memory and superelasticity (recoverable strain up to ~8 %), low modulus; nickel release is the main concern Self-expanding stents, guidewires, orthodontic archwires, bone staples Others Tantalum (porous trabecular metal), gold and silver (dental, antibacterial), platinum and Pt-Ir (electrodes), amalgam (dental), magnesium alloys (biodegradable) Ta — excellent bone ingrowth;

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Pt — inert and conductive, ideal for stimulating electrodes;

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Mg — degrades safely but releases hydrogen Porous implants, electrodes, biodegradable fixation Corrosion and its Control • Corrosion is the electrochemical degradation of a metal: an anodic reaction (M → Mⁿ⁺ + ne⁻) coupled to a cathodic reaction (oxygen reduction in the body).

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The body is an aggressive electrolyte — 0.9 % saline, 37 °C, dissolved oxygen, proteins and cells. • Forms of corrosion in implants: uniform; galvanic (two dissimilar metals in contact — never mix a stainless plate with titanium screws); pitting (localised breakdown of the passive film by chloride); crevice (under a screw head or at a plate-screw interface, where oxygen is depleted and pH falls); intergranular (sensitised stainless steel); fretting (micromotion at a modular taper — the source of much debris in modular hip stems); stress corrosion cracking; and corrosion fatigue. • Protection comes chiefly from passivation — a thin, adherent, self-repairing oxide (Cr₂O₃ on stainless and Co-Cr, TiO₂ on titanium) that isolates the metal.

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Passivation treatment in nitric acid, electropolishing, anodising and coatings all thicken or improve this film.

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Design measures include using a single alloy in a construct, avoiding crevices, and polishing surfaces. • Consequences of ion release: local toxicity and inflammation, metal hypersensitivity (nickel, cobalt and chromium are the common allergens), metallosis and pseudotumour formation around metal-on-metal bearings, systemic distribution, and possible carcinogenicity — the reason for the shift toward nickel-free and vanadium-free alloys.

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Collagen • Collagen is the most abundant protein in the human body (about 25-30 % of total protein) and the principal structural protein of the extracellular matrix, giving tensile strength to skin, tendon, ligament, bone, cartilage, blood vessels and cornea. • Structure: the basic unit, tropocollagen, is a right-handed triple helix of three left-handed polypeptide α-chains, about 300 nm long and 1.5 nm in diameter.

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Its repeating sequence is (Gly-X-Y)n, where every third residue must be glycine (the smallest amino acid, the only one that fits the crowded helix centre) and X and Y are frequently proline and hydroxyproline.

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Hydroxylation of proline and lysine requires vitamin C — its deficiency causes scurvy.

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Molecules assemble into fibrils with a characteristic 67 nm D-banding visible by electron microscopy, stabilised by covalent crosslinks that increase with age. • Types:

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Type I (skin, tendon, bone, the commonest, ~90 % of body collagen), Type II (hyaline cartilage), Type III (reticular fibres, blood vessels, granulation tissue), Type IV (basement membrane, a network rather than fibrils) and Type V. • Advantages as a biomaterial: biodegradable, low immunogenicity (the helical region is poorly antigenic; the non-helical telopeptides carry most of the antigenicity and are removed to make atelocollagen), haemostatic, supports cell adhesion through RGD and other integrin-binding sequences, and is easily processed into films, sponges, gels, tubes and fibres.

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Limitations: poor mechanical strength when reconstituted, rapid resorption, variability, and possible disease transmission (bovine spongiform encephalopathy, hence the use of porcine, ovine or recombinant collagen). • Physical modification: control of fibre orientation and density by extrusion, electrospinning, freeze-drying (which sets pore size by the freezing rate) and moulding; dehydrothermal treatment (DHT) — heating under vacuum to remove water, which forms crosslinks by condensation without adding a chemical; and UV and gamma irradiation, which crosslink but may also fragment the chains. • Chemical modification: crosslinking to slow degradation and raise strength — glutaraldehyde (very effective and the classic treatment for bioprosthetic heart valves, but the residual aldehyde is cytotoxic and promotes calcification), formaldehyde, carbodiimide (EDC/NHS) (a 'zero-length' crosslinker forming direct amide bonds with no residual toxic linker — now preferred), genipin (a natural, far less cytotoxic agent), diisocyanates and epoxides; plus esterification, acylation and grafting of side groups to alter charge, hydrophilicity or drug binding. • Gelatin is denatured (thermally hydrolysed) collagen — cheaper, water-soluble and non-antigenic, but without the triple helix and therefore mechanically much weaker.

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Proteoglycans and Glycosaminoglycans • Glycosaminoglycans (GAGs) are long, unbranched polysaccharides of repeating disaccharide units, in which one sugar is an amino sugar (glucosamine or galactosamine).

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They are strongly negatively charged because of their carboxyl and sulphate groups, and therefore bind large volumes of water, giving the tissue its swelling pressure and resistance to compression — the complement to collagen's tensile strength. • The main GAGs: hyaluronic acid (hyaluronan) — the only one that is non-sulphated and not covalently bound to a protein core, very large, found in synovial fluid, vitreous humour and loose connective tissue, and used clinically as a viscosupplement, in ophthalmic surgery and in dermal fillers; chondroitin sulphate (cartilage, the most abundant); dermatan sulphate (skin, vessels); keratan sulphate (cartilage, cornea); and heparin and heparan sulphate — heparin being the most negatively charged biological molecule and the basis of anticoagulant surface coatings for catheters, oxygenators and stents. • Proteoglycans are GAG chains covalently attached to a core protein, resembling a bottle brush.

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In cartilage, many aggrecan monomers bind through link proteins to a central hyaluronic acid backbone, forming a huge aggregate trapped within the collagen network; the fixed negative charge draws in water and ions, producing the osmotic swelling pressure that lets cartilage carry compressive load and giving it its biphasic, creeping behaviour.

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Other proteoglycans include decorin (regulates collagen fibril diameter), versican and perlecan. • Biomaterials use: hyaluronic acid hydrogels (often crosslinked for durability) for tissue engineering, drug delivery and viscosupplementation; chondroitin sulphate in cartilage scaffolds; heparin coatings for blood-contacting devices; and collagen-GAG copolymer scaffolds — the classic artificial skin (Yannas and Burke) being a collagen-chondroitin-6-sulphate matrix with a controlled pore size and degradation rate.

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This section covers the types, structure, chemistry and properties of the polymers used in medicine, the ceramics and glasses used in medical devices with the classification of bioceramics, and composite biomaterials.

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Fundamentals • A polymer is a long-chain molecule built by repetition of a monomer, made by addition (chain-growth) polymerisation of unsaturated monomers (polyethylene, PMMA, PTFE) or condensation (step-growth) polymerisation with elimination of a small molecule (nylon, polyesters, polyurethanes). • Structure determines properties: linear, branched, crosslinked or network architecture; molecular weight and its distribution (higher MW gives higher strength, as in UHMWPE); crystallinity (crystalline regions give strength, stiffness, opacity and chemical resistance; amorphous regions give toughness and permeability); tacticity; and copolymer arrangement (random, alternating, block, graft). • Thermal behaviour: the glass transition temperature Tg separates the brittle glassy state from the rubbery state, and the melting temperature Tm applies to crystalline regions.

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A polymer used in the body must have the intended behaviour at 37 °C — silicone is rubbery because its Tg is far below body temperature, while PMMA is glassy because its Tg is about 105 °C.

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Thermoplastics soften and can be reshaped on heating; thermosets and elastomers are crosslinked and cannot. • Degradation in the body occurs by hydrolysis (esters, amides — the mechanism exploited in resorbable sutures), oxidation (by inflammatory cell species — the cause of polyurethane and UHMWPE embrittlement), enzymatic attack and environmental stress cracking.

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Sterilisation matters: gamma irradiation crosslinks but also creates free radicals that embrittle UHMWPE over time, ethylene oxide leaves toxic residues requiring aeration, and steam autoclaving destroys polymers with a low Tg or Tm.

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Polymer Key properties Applications UHMWPE (ultra-high molecular weight polyethylene) Very high MW, tough, extremely low friction and good wear resistance; highly crosslinked and vitamin-E stabilised versions reduce wear debris and oxidation Acetabular cups and tibial inserts of joint replacements PMMA (polymethyl methacrylate) Rigid, glassy, transparent, Tg ≈ 105 °C; polymerises in situ with a strongly exothermic reaction and shrinkage; residual monomer is toxic Bone cement, intraocular lenses, hard contact lenses, dentures Silicone (polydimeth ylsiloxane) Si-O backbone, very flexible, excellent thermal, oxidative and biological stability, hydrophobic, gas-permeable Breast implants, catheters, tubing, drains, shunts, finger joints, membrane oxygenators Polyurethanes Segmented block copolymers with hard and soft segments — tough, elastic, good blood compatibility; polyester types hydrolyse, polyether types oxidise Vascular grafts, catheters, heart-assist diaphragms, wound dressings, pacemaker lead insulation PTFE / ePTFE (Teflon) Extremely inert and hydrophobic, very low friction; expanded PTFE is porous and allows tissue ingrowth Vascular grafts, hernia mesh, sutures, catheter liners PET (Dacron) Strong polyester fibre, knitted or woven Large-diameter vascular grafts, suture, heart-valve sewing rings PVC Cheap, clear, flexible when plasticised — but the plasticiser (DEHP) leaches and is a regulatory concern Blood bags, tubing, dialysis sets Polymer Key properties Applications PGA, PLA, PLGA, PCL, PDO Resorbable aliphatic polyesters degrading by hydrolysis; rate tuned by composition and crystallinity (PGA fastest, PCL slowest); acidic degradation products Resorbable sutures, fixation pins and screws, drug-delivery microspheres, tissue-engineering scaffolds Hydrogels (PHEMA, PVA, PEG, alginate, chitosan) Crosslinked hydrophilic networks holding large amounts of water; soft and tissue-like;

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PEG resists protein adsorption Soft contact lenses, wound dressings, drug delivery, cell encapsulation, antifouling coatings PEEK Semi-crystalline, radiolucent, modulus close to cortical bone, chemically stable and steam-sterilisable Spinal cages, trauma plates, dental abutments Ceramics and Glasses • Ceramics are inorganic, non-metallic materials with strong ionic and covalent bonding, which explains every characteristic property: very high hardness, high melting point, chemical inertness, high compressive strength and excellent wear resistance, but no plastic deformation, low tensile strength and low fracture toughness — they are brittle and flaw-sensitive, so strength is governed by the largest defect and is described statistically by the Weibull distribution.

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Glasses are the amorphous counterparts, based on a silica network with modifiers. • Processing is by powder pressing and sintering, sol-gel or melt processing; porosity and grain size control strength, and surface finish is critical for a bearing.

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Class of bioceramic Behaviour Examples and uses Bioinert Almost no chemical interaction; a thin fibrous capsule forms; relies on mechanical fixation Alumina (Al₂O₃) — very hard, excellent wear, femoral heads; zirconia (ZrO₂, yttria-stabilised) — tougher than alumina by transformation toughening but liable to low-temperature degradation; pyrolytic carbon — outstanding blood compatibility, used in mechanical heart valve leaflets Bioactive (surface reactive) Form a carbonated hydroxyapatite layer in body fluid and bond chemically to bone Bioglass 45S5 (45 % SiO₂, 24.5 % CaO, 24.5 % Na₂O, 6 % P₂O₅) — the original bone-bonding glass; hydroxyapatite (HA); glass-ceramics such as A-W; used as coatings, granules and middle-ear implants Bioresorbable Dissolve progressively and are replaced by new bone Tricalcium phosphate (β-TCP), calcium sulphate, biphasic HA/TCP, calcium phosphate cements — bone graft substitutes and void fillers • Hydroxyapatite, Ca₁₀(PO₄)₆(OH)₂, with a Ca/P molar ratio of 1.67, is the mineral phase of bone and teeth (biological apatite being carbonate-substituted, non-stoichiometric and nanocrystalline).

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It is osteoconductive and bioactive but brittle and weak in tension, so it is used chiefly as a plasma-sprayed coating on titanium implants, as granules and as the reinforcing phase of composites. β-TCP has a Ca/P ratio of 1.5 and resorbs considerably faster. • Carbon materials deserve separate note: pyrolytic carbon is the standard material for mechanical heart valve components because of its combination of thromboresistance, wear resistance and fatigue strength; diamond-like carbon is used as a thin hard coating.

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Composites • A composite combines a matrix (polymer, metal or ceramic) with a reinforcement (particles, short fibres, continuous fibres or a second network) to obtain properties that neither phase has alone.

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The properties depend on the volume fraction, aspect ratio and orientation of the reinforcement and, decisively, on the quality of the interface, through which load is transferred. • The rule of mixtures gives the upper (iso-strain, fibres aligned with the load) bound Ec = EfVf + EmVm and the lower (iso-stress, transverse) bound 1/Ec = Vf/Ef + Vm/Em.

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Aligned fibre composites are therefore strongly anisotropic — like bone itself. • Natural composites are the model: bone (hydroxyapatite in a collagen matrix — a ceramic-polymer composite) and cartilage (proteoglycan gel in a collagen network). • Medical composites: dental restorative resins — silica or glass filler (typically 60-80 % by volume) in a Bis-GMA or UDMA matrix, light-cured, where a high filler content raises strength and lowers the polymerisation shrinkage that causes marginal leakage;

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HA-reinforced polyethylene (HAPEX), designed to match the modulus of bone; carbon-fibre-reinforced PEEK and epoxy for plates, spinal implants and prosthetic limbs; fibre-reinforced bone cement; and collagen-GAG and polymer-ceramic scaffolds for tissue engineering. • Advantages: modulus can be matched to bone, avoiding stress shielding; high strength-to-weight ratio; radiolucency; and tailored degradation.

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Limitations: complex and expensive manufacture, interfacial debonding and delamination, difficulty of sterilisation and quality control, and release of particulate debris if the matrix wears.

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This section covers the general principles of surface modification, the methods used to modify surfaces for enhanced biological interaction, the nature and production of the plasma environment, and high-energy and high-temperature plasma treatments.

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General Principles of Surface Modification • The governing idea is that the bulk provides the mechanical performance while the surface determines the biological response.

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Surface modification therefore aims to keep the bulk properties unchanged while altering only the outermost layer, and it is used to make a surface more or less wettable, to change its charge or chemistry, to add a bioactive or drug-releasing layer, to improve wear or corrosion resistance, or to prevent protein, cell or bacterial adhesion. • Principles of good practice: the modified layer should be thin (so as not to alter bulk behaviour), strongly adherent (delamination produces debris and is a common failure mode), uniform and reproducible, stable against remodelling, leaching, wear and sterilisation, and free of contamination — surfaces must be rigorously cleaned first, since biological performance can be destroyed by a monolayer of silicone or mould release. • Classification: modification may be physicochemical (altering the existing surface — plasma, etching, oxidation, ion implantation, laser or mechanical texturing), by coating (adding a distinct layer — plasma spraying, PVD, CVD, dip and spin coating, electrodeposition, sol-gel, Langmuir-Blodgett films, self-assembled monolayers), or biological (grafting or immobilising molecules — covered in 2.5).

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Methods for Modifying Surfaces to Enhance Biological Interaction Method Principle and effect Mechanical texturing (grit blasting, machining, polishing) Roughness controls cell response: rough, microtextured surfaces favour bone cell attachment and osseointegration, while smooth, polished surfaces are preferred for blood contact and for bearing surfaces Chemical treatment (acid etching, alkali treatment, oxidation, passivation) Cleans, creates micro-roughness and functional groups; alkali-heat treatment of titanium produces a bioactive sodium titanate layer Anodisation Electrochemical thickening of the TiO₂ film, giving controlled thickness, colour and nanotube arrays that enhance osteoblast response and can carry drugs Plasma spraying Molten particles projected onto the substrate — the standard industrial route for hydroxyapatite coatings on titanium hip stems and dental implants; thick (50-200 μm), but with a risk of phase change and delamination PVD (sputtering, evaporation, arc) Atoms ejected from a target condense as a very thin, dense, adherent film — TiN, DLC, hard wear coatings CVD / PECVD Film grown from gaseous precursors reacting at the surface; plasma-enhanced CVD allows deposition at low temperature, essential for polymers Ion implantation and ion beam treatment High-energy ions driven into the surface, changing composition and hardness with no distinct interface and therefore no delamination — nitrogen implantation hardens titanium bearings Sol-gel, dip, spin and spray coating Simple, low temperature, conformal; used for HA, silica, drug-eluting and polymer layers Self-assembled monolayers (SAMs) and Langmuir-Blodgett films Ordered molecular monolayers (alkanethiols on gold, silanes on oxides) giving precisely defined surface chemistry — the standard research tool for studying protein and cell response Grafting and plasma polymerisation Covalent attachment of polymer chains — PEG grafting to resist protein adsorption, heparin grafting for anticoagulation, charged groups for cell adhesion Drug-eluting and antibacterial coatings Polymer reservoir releasing an antiproliferative agent (drug-eluting stents), an antibiotic, or silver ions • Silane coupling agents deserve special mention as the standard bridge between an inorganic oxide surface and an organic polymer or biomolecule: one end hydrolyses and condenses with surface hydroxyls, the other carries an amine, epoxy or methacrylate group for the organic phase — the same chemistry used in dental composites and in immobilising proteins on glass (2.5).

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The Nature and Production of the Plasma Environment • A plasma is the fourth state of matter: a partially ionised gas containing free electrons, ions, radicals, excited neutrals and photons, which is electrically conductive but macroscopically neutral (quasi-neutral) because the densities of positive and negative charges are equal. • Production: a gas at reduced pressure (typically 0.1-10 mbar) is subjected to an electric field — DC, radio-frequency (usually 13.56 MHz) or microwave (2.45 GHz) — which accelerates stray electrons until they ionise gas molecules by collision, producing further electrons in an avalanche sustained as a glow discharge.

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Atmospheric-pressure plasmas (corona, dielectric barrier discharge, plasma jet) avoid the vacuum system and can treat heat-sensitive and large objects. • The crucial distinction — thermal versus non-thermal plasma: in a non-thermal (cold, non-equilibrium, low-temperature) plasma, the electrons are extremely energetic (equivalent to tens of thousands of kelvin, 1-10 eV) while the heavy ions and neutrals remain near room temperature, because the light electrons gain energy from the field but transfer little of it in collisions with heavy particles.

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The gas therefore stays cool while the surface chemistry is highly energetic — which is precisely what allows plasma treatment of polymers and even biological materials without thermal damage.

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In a thermal (hot, equilibrium) plasma — an arc or a plasma torch — all species share a high temperature of thousands of kelvin. • Process gases and their effects: argon and helium (inert — physical etching, cleaning, crosslinking by radical formation, the so-called CASING process), oxygen (oxidation, introduction of polar -OH, -C=O and -COOH groups, strong increase in wettability, ashing of organic contamination), nitrogen and ammonia (amine groups for subsequent coupling of biomolecules), fluorine-containing gases such as CF₄ (fluorination, making the surface strongly hydrophobic and lowering friction), and monomer vapours for plasma polymerisation, which deposits a thin, pinhole-free, highly crosslinked and strongly adherent film of almost any chemistry on almost any substrate.

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High-Energy and High-Temperature Plasma Treatments • Low-temperature (cold) plasma treatment is the workhorse for medical polymers: cleaning and sterilisation (low-temperature hydrogen peroxide plasma sterilisation is standard for heat-sensitive devices), surface activation to improve adhesion, printing and bonding, wettability control, etching, crosslinking, plasma polymerisation and plasma grafting, which creates surface radicals to which monomers such as acrylic acid or PEG are then covalently attached. • High-energy treatments: plasma immersion ion implantation (PIII) and beam-line ion implantation drive energetic ions tens of nanometres into the surface, producing a graded, non-delaminating modified layer that hardens the surface and improves wear and corrosion resistance; ion beam assisted deposition (IBAD) uses simultaneous bombardment to densify a growing film and improve its adhesion; magnetron sputtering gives dense thin films of TiN, DLC and HA. • High-temperature (thermal) plasma treatments: the plasma spray torch, in which powder is injected into an arc plasma at several thousand kelvin, melted and accelerated onto the substrate.

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This is how hydroxyapatite is coated onto titanium implants, and how titanium and porous metal layers are applied.

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The advantages are speed, thickness and line-of-sight versatility; the drawbacks are thermal decomposition of HA into other calcium phosphates and amorphous phases, residual stress, variable crystallinity and the risk of coating delamination — hence the interest in lower-temperature alternatives such as sol-gel, biomimetic deposition in simulated body fluid, and electrochemical deposition. • Limitations of plasma processing generally: it is line-of-sight and treats the inside of narrow lumens poorly; the effect is not permanent — treated polymer surfaces undergo hydrophobic recovery (ageing) as mobile chains rotate the new polar groups into the bulk, so devices must be used or coated soon after treatment; vacuum systems are costly; and the process must be validated for each geometry and material.

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This section covers biologically active molecules, the solid supports used to carry them, the methods of immobilizing them, and the applications of biologically functional materials.

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Biologically Active Molecules • A biologically functional (bioactive) material is one that carries a biological molecule at its surface so that the device communicates with the tissue in biological language rather than merely tolerating it.

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Instead of asking a synthetic surface to be ignored, the molecule tells the adjacent cells what to do. • Classes of molecule used: enzymes (glucose oxidase, urease, heparinase — the recognition element of most biosensors); antibodies and antigens (immunoassays, affinity separation, targeting); peptides and proteins — especially the cell-adhesion sequence RGD (arginine-glycine-aspartic acid) from fibronectin, which binds cell integrins and is the single most-used bioactive motif, along with laminin-derived YIGSR and IKVAV; growth factors (BMP-2 for bone, VEGF for vessels, EGF and FGF); extracellular matrix proteins (collagen, fibronectin, laminin, vitronectin); nucleic acids and aptamers (DNA microarrays, gene delivery, biosensing); polysaccharides (heparin for anticoagulation, hyaluronic acid, chitosan); drugs and antibiotics; and whole cells or micro-organisms. • Why immobilise rather than release? Immobilisation gives reuse and long service life, keeps the molecule localised where it is needed, avoids systemic side effects, stabilises the molecule against denaturation and proteolysis, allows the product to be free of the biomolecule, and permits continuous rather than batch operation.

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The price is possible loss of activity, diffusional limitation and altered kinetics.

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Solid Supports for Biomolecules Support Features Inorganic — glass, silica, controlled-pore glass, alumina, titania Rigid, dimensionally stable, resistant to microbial attack and solvents, easily silanised; but brittle and with limited functional groups Synthetic polymers — polystyrene, PMMA, polyacrylamide, nylon, polyurethane, PVDF membranes Wide range of chemistry and form, cheap; polystyrene microplates are the standard ELISA support Natural polymers — agarose, cellulose, dextran (Sephadex), alginate, chitosan, collagen Highly hydrophilic and biocompatible, mild coupling conditions, good retention of activity; but mechanically weak and biodegradable Metals and carbon — gold, platinum, carbon, graphene Electrically conductive, essential for electrochemical biosensor electrodes; gold-thiol chemistry gives excellent self-assembled monolayers Magnetic particles Allow rapid magnetic separation and recovery — the basis of modern immunoassay and cell-sorting platforms Hydrogels and membranes High water content and permeability, mild environment for entrapped cells and enzymes • Requirements of a good support: high surface area and appropriate porosity, suitable functional groups for coupling, mechanical and chemical stability, hydrophilicity, insolubility, non-toxicity, resistance to microbial attack, low non-specific adsorption, and acceptable cost and regenerability.

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Immobilization Methods Method Principle Strengths and weaknesses Physical adsorption Van der Waals, hydrophobic and electrostatic binding to the support — simply incubate and rinse Simplest, cheapest, mildest, no reagents and little loss of activity; but weak, and the molecule leaches with changes of pH, ionic strength or temperature, and may be displaced by other proteins Ionic (electrostatic) binding Binding to an oppositely charged ion-exchange support (DEAE-, CM-cellulose) Mild and reversible, so the support can be regenerated; but sensitive to pH and salt concentration Covalent bonding Formation of a stable covalent bond between functional groups of the biomolecule (-NH₂ of lysine, -COOH, -SH of cysteine, -OH) and activated groups on the support.

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Typical chemistries: glutaraldehyde (amine to amine), carbodiimide EDC/NHS (carboxyl to amine, zero-length), cyanogen bromide on polysaccharides, epoxy, tosyl and maleimide (to thiols), and silanisation followed by coupling on oxide surfaces Strongest and most stable, no leaching, and the orientation can be controlled; but harsh conditions may denature the molecule or block the active site, it is irreversible, and the support cannot be reused Crosslinking Intermolecular linking of biomolecules to each other with a bifunctional reagent (glutaraldehyde), with or without an inert protein such as albumin, forming insoluble aggregates or CLEAs No support needed and very stable; but often substantial activity loss and poor mechanical properties Entrapment Physical confinement within a polymer gel or fibre network — calcium alginate beads, polyacrylamide, sol-gel silica, PEG hydrogels Mild, widely applicable, and the molecule is not chemically altered; but diffusion limitation, possible leakage of small molecules and a lower effective activity Microencapsulat ion Enclosure inside a semipermeable membrane capsule which admits substrates and releases products but retains the biomolecule or cell Large loading and protection from the immune system (the principle of encapsulated islet cells for an artificial pancreas); but membrane fouling and rupture Affinity immobilisation Specific biological recognition — biotin-avidin/streptavidin (the strongest known non-covalent interaction), protein A to the Fc of antibodies, His-tag to nickel Excellent, uniform orientation and retained activity, which is why it dominates modern biosensor and microarray fabrication; but costly and requires prior modification • Choosing a method is a compromise between stability (covalent, crosslinking) and retained activity (adsorption, entrapment, affinity).

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The key design variables are the orientation of the molecule (an antibody must present its Fab sites outward; an enzyme's active site must not be blocked or attached through), the surface density and spacing (a spacer arm such as a PEG chain holds the molecule away from the surface and restores much of its native mobility and activity), and the micro-environment (charge and hydrophilicity of the support shift the apparent optimum pH and Km).

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Applications • Biosensors — an immobilised bioreceptor (enzyme, antibody, DNA, cell) on a transducer (electrochemical, optical, piezoelectric, thermal).

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The classic example is the glucose biosensor, with glucose oxidase immobilised on an electrode, which is the basis of the entire self-monitoring and continuous glucose monitoring industry; these are taken further in the instrumentation chapters. • Diagnostics — ELISA plates with adsorbed antibody, lateral-flow immunoassays, DNA and protein microarrays, immunomagnetic cell separation. • Blood-contacting devices — heparinised surfaces on catheters, oxygenators, dialysers and stents; surfaces bearing thrombomodulin, tPA or nitric-oxide-releasing groups; and albumin pre-coating to reduce platelet adhesion. • Tissue engineering and implants — RGD-grafted scaffolds to promote specific cell adhesion, growth-factor-loaded scaffolds (BMP-2 on collagen for spinal fusion), and cell-instructive surfaces that direct stem-cell differentiation. • Antibacterial surfaces — immobilised antibiotics, antimicrobial peptides, quaternary ammonium compounds, lysozyme and silver. • Bioreactors and industrial/biomedical processing — immobilised enzymes for continuous conversion, immobilised urease in artificial kidney sorbents, affinity columns for plasma purification (LDL apheresis, endotoxin removal), and encapsulated cells for the bioartificial liver and pancreas. • Drug delivery — molecules bound to or entrapped in a carrier for controlled or targeted release, including antibody-targeted nanoparticles.

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This section covers inflammation, wound healing and the foreign body response, the interaction of endothelial cells and the extracellular matrix with biomaterials, blood-biomaterial interaction, and bacteria and biomaterials.

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The Sequence of Events at an Implant Surface Implantation injures tissue, and the response follows a fixed sequence whose timescales are worth memorising: protein adsorption (seconds) → blood-material interaction and provisional matrix formation (minutes to hours) → acute inflammation (hours to days) → chronic inflammation (days to weeks) → granulation tissue and foreign body reaction (weeks) → fibrous encapsulation (weeks to months). • Protein adsorption is the very first event and the most important: within seconds the surface is covered by a layer of plasma proteins, so cells never encounter the bare material — they encounter the adsorbed protein layer, and its composition and conformation determine everything that follows.

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The Vroman effect describes the sequence: abundant, highly mobile proteins (albumin) adsorb first and are progressively displaced by less abundant proteins of higher surface affinity (fibrinogen, then high-molecular-weight kininogen, fibronectin, factor XII).

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Adsorbed proteins may also unfold (change conformation), exposing hidden epitopes that activate platelets and complement — which is why an adsorbed protein is not the same as the protein in solution. • Surface properties that govern adsorption: hydrophobic surfaces adsorb more protein and denature it more; charged surfaces bind proteins of opposite charge; rough surfaces present more area; and highly hydrated, neutral, mobile surfaces such as PEG resist adsorption, which is the basis of every antifouling strategy.

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Inflammation and Wound Healing • Acute inflammation (minutes to days) is the immediate vascular and cellular response to injury, with the five cardinal signs rubor (redness), calor (heat), tumor (swelling), dolor (pain) and functio laesa (loss of function).

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Vasodilatation and increased permeability produce exudate, and neutrophils arrive first, followed by monocytes that become macrophages.

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Mediators include histamine, prostaglandins, kinins, complement and cytokines (IL-1, IL-6, TNF-α). • Chronic inflammation (beyond about two weeks) is characterised by macrophages, lymphocytes and plasma cells rather than neutrophils, and by simultaneous tissue destruction and repair.

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Persisting chronic inflammation around an implant is a sign of poor biocompatibility, of motion, of wear debris or of infection. • Wound healing phases: haemostasis (clot and provisional fibrin matrix) → inflammation → proliferation (granulation tissue — fibroblasts laying down type III collagen, angiogenesis, re-epithelialisation, wound contraction by myofibroblasts) → remodelling/maturation (months to years; type III collagen replaced by type I, tensile strength recovering only to about 70-80 % of the original). • Healing may be by primary intention (clean, apposed edges — minimal scar), secondary intention (open wound with tissue loss, granulation and contraction — much more scar) or tertiary/delayed primary intention.

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It is impaired by infection, ischaemia, diabetes, corticosteroids, malnutrition, foreign bodies and irradiation.

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The Foreign Body Response • When the implant cannot be phagocytosed because it is too large, macrophages attempt frustrated phagocytosis, releasing reactive oxygen species and degradative enzymes onto the surface, and then fuse with one another to form multinucleated foreign body giant cells (FBGCs) — the histological hallmark of the foreign body reaction.

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The end result is a fibrous (collagenous) capsule that walls the device off from the tissue. • Interpretation: a thin capsule (tens of micrometres) with few giant cells indicates good biocompatibility; a thick capsule with persistent giant cells and chronic inflammation indicates a poor response.

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The capsule is harmless around some devices but disastrous around others — it prevents osseointegration, it causes capsular contracture around breast implants, it insulates electrodes and raises stimulation thresholds, and it blocks the diffusion of glucose to an implanted sensor, which is the principal reason implantable biosensors drift and fail. • Factors that intensify the reaction: particulate and irregular shapes and rough surfaces (particles and fibres provoke far more reaction than a smooth monolith of the same material), micromotion, wear debris (UHMWPE particles in the 0.1-10 μm range are the most biologically active and drive osteolysis and aseptic loosening — the leading cause of late joint-replacement failure), leachables and degradation products, and infection.

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Endothelial Cells, ECM and Biomaterial Interaction • The extracellular matrix is the natural biomaterial: a network of structural proteins (collagen, elastin), adhesion proteins (fibronectin, laminin, vitronectin) and proteoglycans and GAGs, which provides mechanical support, anchors cells, stores growth factors and instructs cell behaviour.

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Cells attach to it through integrins — transmembrane receptors that bind RGD and similar motifs and cluster into focal adhesions linked to the actin cytoskeleton.

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Through these, cells sense not only chemistry but also substrate stiffness and topography (mechanotransduction), which is why matrix stiffness can direct stem-cell fate and why surface texture alters cell response. • Endothelial cells line every blood vessel and form the only truly non-thrombogenic surface known.

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A healthy endothelium is actively antithrombotic — it presents a negatively charged glycocalyx and secretes prostacyclin (PGI₂) and nitric oxide (which inhibit platelet aggregation and cause vasodilatation), thrombomodulin (which switches thrombin to activating protein C), heparan sulphate and tissue plasminogen activator.

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When it is damaged or activated it becomes prothrombotic, exposing collagen and von Willebrand factor and expressing tissue factor and adhesion molecules. • Endothelialisation of devices is therefore the ideal solution to blood compatibility: in vitro seeding of vascular grafts with autologous endothelial cells, in situ capture of endothelial progenitor cells with anti-CD34 antibody coatings (the basis of some endothelial progenitor cell capture stents), and surface modification with RGD, fibronectin, VEGF or nitric-oxide-releasing chemistry to encourage it.

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The difficulty is that small-diameter (< 6 mm) synthetic grafts still fail through thrombosis and intimal hyperplasia because a stable, confluent, correctly orientated endothelium is hard to achieve and maintain under flow — the outstanding unsolved problem of vascular device design.

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Blood-Biomaterial Interaction • A foreign surface in contact with blood triggers four interlocking cascades: protein adsorption → platelet adhesion, activation and aggregation → activation of the intrinsic coagulation cascade → complement and leukocyte activation, ending in thrombus. • Protein layer: adsorbed fibrinogen is strongly platelet-adhesive (its exposed sequences bind the platelet GPIIb/IIIa receptor), whereas adsorbed albumin is passivating — which is why albumin pre-coating reduces platelet adhesion and why the fibrinogen-to-albumin ratio on a surface predicts its thrombogenicity. • Platelets adhere (via vWF and GPIb), change shape and spread, release ADP, thromboxane A₂ and serotonin from their granules, and aggregate; activated platelet membranes then provide the phospholipid surface on which the coagulation complexes assemble. • Coagulation: as noted in Chapter 1, a foreign surface activates factor XII and the intrinsic (contact) pathway, proceeding to thrombin and fibrin. • Complement is activated mainly by the alternative pathway on surfaces bearing hydroxyl and amine groups — cellulose dialysis membranes are the classic example, producing the anaphylatoxins C3a and C5a, leukocyte activation and the transient neutropenia and hypoxaemia of first-use dialysis reactions.

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This is why cellulose membranes have largely been replaced by synthetic polysulphone and polyacrylonitrile membranes. • Haemolysis and cell damage from high shear must also be considered, as in 1.6. • Design strategies for blood compatibility: smooth, defect-free surfaces with no stagnation or recirculation zones and no excessive shear; hydrophilic, protein-resistant coatings (PEG, phosphorylcholine — which mimics the outer leaflet of the cell membrane and is used on stents and circuits); bioactive coatings (heparin, thrombomodulin, nitric oxide donors); negatively charged surfaces, since platelets and the endothelium are themselves negatively charged; endothelialisation; and, failing all these, systemic anticoagulation, which remains necessary for mechanical heart valves and extracorporeal circuits.

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Bacteria and Biomaterials • Device-related infection is among the most serious complications in implant surgery.

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Its central concept is the race for the surface: if host tissue cells colonise the implant first, the surface is defended; if bacteria arrive first, they establish a biofilm that tissue cannot displace. • Biofilm formation: reversible adhesion → irreversible attachment via adhesins → microcolony formation and secretion of an extracellular polymeric substance (EPS, glycocalyx) → maturation into a structured three-dimensional biofilm → dispersal of cells that seed new sites. • Why biofilms are so difficult to treat: the EPS matrix is a diffusion barrier to antibiotics and to immune cells and antibodies; cells deep in the biofilm are metabolically dormant (persister cells) and most antibiotics act only on dividing organisms; quorum sensing coordinates virulence; and gene transfer spreads resistance.

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Bacteria in a biofilm may tolerate 100 to 1,000 times the antibiotic concentration that kills the same organism when free-floating (planktonic), which is why an infected implant usually has to be removed. • Common organisms:

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Staphylococcus epidermidis (the classic coagulase-negative biofilm former on catheters, shunts and prostheses), Staphylococcus aureus (including MRSA), Pseudomonas aeruginosa, enterococci and Candida.

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The presence of a foreign body lowers the infecting dose enormously — a well-known experiment showed that the number of staphylococci needed to cause infection falls by several orders of magnitude when a suture is present. • Prevention strategies: strict asepsis, sterilisation and short operating time; antibiotic prophylaxis and antibiotic-loaded bone cement or beads; anti-adhesive surfaces (PEG, phosphorylcholine, very smooth finishes, low surface energy); antibacterial surfaces — silver ions and nanoparticles, copper, quaternary ammonium compounds, chlorhexidine, antimicrobial peptides, nitric oxide release, and contact-killing nanostructured (nanospike) surfaces; drug-eluting coatings; and design measures avoiding crevices and dead spaces.

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Treatment of an established infection generally requires debridement and device removal or exchange with prolonged antibiotics.