The UHMWPE powder coming from the Ziegler-Natta polymerisation plant is processed by compression moulding and ram extrusion: both techniques use high pressure and controlled heating and cooling cycles, and do not significantly modify chemical, physical and structural characteristics of the starting polymer, with the exception of crystallinity (which is normally much higher in the pristine powder). Therefore all prosthetic components, ready to be sterilised, still retain all properties of the starting material.

The main sterilisation processes used nowadays employ ethylene oxide (EtO), gas-plasma (GP) and high-energy radiation (gamma radiation and electron beam) [9–11].

EtO and GP are surface sterilization methods and do not significantly affect the physical, chemical and mechanical properties of prosthetic components. GP is based on the action of ionized gas (i.e. hydrogen peroxide or peracetic acid).

Gamma radiations are emitted during decay of a ⁶⁰Co unstable nucleus. The dose absorbed by prosthetic components is about 25–30 kGy and depends upon the geometry of the sample and its position in relation with the source.

Electron beam is produced by thermally exciting a tungsten filament; the emitted electrons are accelerated by electric fields up to 10 MeV and then conveyed onto the material to be sterilised. The advantages of this method are the easy control of the apparatus and the very short period of treatment (seconds).

Gamma radiation and electron beam have a mean energy some orders of magnitude higher than that of polymeric chemical bonds and therefore generate the scission of some chemical bonds of the UHMWPE and formation of free radicals. If even a single C-C bond of the UHMWPE chain is broken and 2° CH₂ ~ radicals are formed, the length of the chain and consequently the molecular mass decrease, with worsening of some chemical and physical material characteristics. This process is called degradation and in presence of oxygen, oxidation, which involves free radicals (Fig. 5.1).

The oxidative process depends on the radicals (formed during sterilisation) and on the amount of oxygen diffused into the PE components from the atmosphere during processing, sterilisation if conducted in presence of air and storage [12].

The distribution of oxidative products in the prosthetic component depends from the following variables: rate at which radiations is supplied, temperature of the sterilisation chamber, amount of oxygen present in the polymer when irradiated and diffused afterwards. Both in new and retrieved component, a crown effect or white band was the macroscopic evidence of this oxidation, responsible for many severe failures (delamination and fracture) during service in vivo in years ‘90’. Unfortunately, the first dramatic failures of UHMWPE components in the mid 1980s were attributed to inadequate mechanical properties of the UHMWPE, despite the evidence that these properties were much better than those required by ASTM F648.

An adequate packaging of the components is mandatory to assure the correct atmosphere in accordance with the chosen sterilization process; the packaging could be critical when high energy radiation in vacuum or inert gases to reduce oxidation is used. Currently employed packaging can be included in three categories [13]:

Ultimately, a complete absence of oxidation is obtained only by gas-sterilisation.

Polyethylene debris are particles loss due to friction, caused by the reciprocal movement of the loaded articular surfaces: for equal mechanical stress, material and interface, abrasion is function of time. Whereas dramatic failures due to anomalous wear of heavily oxidised polyethylene have become quite uncommon nowadays, the production of abraded particles remains a problem in young patients whose life expectancy and quality of life are very high. The debris initiate an inflammatory reaction, the formation of a loosening membrane and a secondary osteolysis. The junctional tissue depends from number, size and chemical structures of UHMWPE debris. While pointing out that this topic is in continuous development, it is important to realise that the debris is not just simple UHMWPE particles, but biologically active particles whose surface interact with the human tissues according with their macro and micromorphology, contact area, molecules adsorpted on their surface, superficial hydrophilic and hydrophobic character, release of free radicals and time [9–11].

A process of adsorption and deep diffusion into the UHMWPE prosthetic components of organic molecules present in the synovial liquid, such as cholesterol, ester of cholesterol, squalene, β-carotene, takes place in vivo. This diffusion explains the yellowish colour in some retrieved components [14].

To increase the abrasion resistance, crosslinked UHMWPE (X-PE) appeared on the market in the late 1990s [9–11, 15]. Crosslinking of a polymer is the linking of two or more molecular chains by means of chemical covalent bonds: macro radical species, formed by treatment with high energy, react with vinyl double bonds, linking the polymer chains with a C-C stable chemical bond and giving Y-crosslink. The X-PE can be represented as one long, branched molecule with infinite molecular mass and consequent better wear resistance properties than standard UHMWPE, but also with some lower mechanical properties, owing to chemical and physical modifications induced by irradiation and heat treatment.

Commercially available X-PEs are obtained by different crosslinking processes, mainly based on gamma radiation or electron beam at doses ranging from 60 to 100 kGy at room temperature or in the molten state, depending on the manufacturer; the residual radicals are eliminated by thermal treatment, sometime at temperature below the melting point of the polymer (typically at 130 °C) (annealing). The final sterilization is obtained by EtO or gas-plasma or, in few cases, by gamma radiation in low oxygen environment [12].

Due to different crosslinking processes, the commercial X-PEs can be very different with variable properties, while standard UHMWPE has and maintain its properties if processed and sterilised in adequate ways.

Even if dramatic oxidation levels are not observed in newly produced UHMWPE components, it must be kept in mind that also very low oxidation levels can lead to significant variations in the mechanical properties of the polymer.

Vitamin E or, better, its synthetic derivative, alfa-tocopherol, is employed to stabilize UHMWPE against oxidation (ASTM F2695-12). As already pointed out, PE is easily subject to oxidation, which strongly compromises their mechanical properties. The oxidation is basically due to the reaction between macroradicals and oxygen diffused into the polymer from the surrounding atmosphere; Vitamin E decreases the macro alkyl radicals available to react with the oxygen and thus to a significant slowdown of the oxidative cascade [9–11, 15–17]. Unfortunately, a decreased number of available alkyl radicals is also responsible for a lower efficiency of crosslinking at the same radiation dose, but a correct vitamin E concentration and radiation dose determine an oxidatively stable UHMWPE, without the need of a further thermal treatment, with enough crosslink density and consequent resistance to abrasion.

Orthopaedic cement is basically poly(methyl methacrylate) (PMMA) obtained by polymerising the methyl methacrylate monomer (MMA) [18, 19]. Usually it is supplied in two separate packages: a brown coloured vial (in order to avoid any negative influence of the light on the monomer) containing about 20 ml of transparent liquid, and one package or two containing 40 g of powder. The liquid contains: MMA, usually N,N dimethyl-p-toluidine (DMPT) to accelerate the polymerisation process in presence of radicals, and traces of hydroquinone to avoid premature polymerisation of the monomer. The powder is formed by pre-synthesised PMMA (at times polymethylmethacrylate-styrene as copolymers are used), dibenzoyl peroxide (DBP) and barium sulphate (or zirconium dioxide), the latter may be supplied in a separate package. PMMA is in the shape of spherical particles having a variable diameter between 30 and 250 μm; the size of the particles determines the viscosity of the cement. When the contents of the two packages are mixed, DBP initiates the radical process of polymerisation through polymerisation accelerator and the effect of polymerisation heat. Barium sulphate makes the cement radio-opaque.

Cements produced by different industrial companies have different chemical-physical characteristics and mechanical properties due various components and their relative concentrations.

Bone cement preparation is characterised by three phases: the wetting phase corresponds to mixing the solid part with the liquid, the setting phase (divided into ‘dough time’ and ‘working time’) corresponds to the initial polymerisation process (about 5 % of total), the curing phase corresponds to the final hardening phase and completion of the polymerisation process. During mixing, benzoyl peroxide, present on the surface of the PMMA powder, and DMPT present in the liquid, interact and the polymerisation process starts, mainly on the surface of the pre-synthesised poly(methyl methacrylate). Working time starts when a “dough” is obtained which no longer sticks to gloves and temperature increase of the cement is minimal, corresponding to minimal transformation of MMA to PMMA. The final polymerisation phase is characterised by the rapid increase of polymerisation rate and temperature. The time required for the various phases depend mainly on the temperature in the operating theatre: a 10 °C increase causes polymerisation to start twice as quickly, cutting mixing times by half. After polymerization, less than 5 % of MMA remains free and this percentage may slowly spread into the body. The MMA polymerization reaction is exothermic; the high temperature favours DBP decomposition leading to an increase in radical formation and consequently an increase in polymerization process. Therefore, polymerization speed is initially minimal and gradually increases. Where processing carried out in adiabatic conditions, the bone cement temperature would reach 160 °C. The actual temperature reached by the cement during the surgery depends on the balance between quantity and speed with which the heat is produced, and how easily the heat is dispersed from the surface into surrounding tissues. At the interface with spongy bone, due to vascularisation and the trabecular shape of the bone itself, temperatures of 60 °C can be reached, while in the centre of the mass of cement the temperature is higher than 100 °C. Schematically cement produces heat in function of the used amount, and the temperature at the interface increases with the higher quantity of cement. Based on this assumption, an adequate surgical technique can lower the temperature at the interface by using both an adequate and not too thick layer of cement, and washing liquids in the final polymerization phase. Some cements are declared as “low temperature polymerization”. They are characterised by a lower ratio monomer MMA/polymer that proportionally lowers the heat developed during transformation of monomer into polymer. High temperature is sought when the cement is used as adjuvant in bone tumours to ensure “sterilisation” of a bone surface from which the tumour has been removed; therefore in oncological surgery, standard PMMA is useful.

During polymerization reaction, a theoretical volumetric shrinking of the PMMA takes place proportional to the amount of MMA used; in the orthopaedic cement, the volumetric shrinking is 7 % of the initial volume. Another characteristic of cement is the porosity due to CO₂ formed during decomposition of the initiator, MMA monomer evaporation, air-bubble formed during hand preparation of the mixture, and the expansion due to temperature increase during polymerisation. In actual orthopaedic cements, the vacuum technique preparation decreases air-bubble formation; other factors cannot be eliminated.

Antibiotic-loaded cements are used in order to obtain a greater quantity of local antibiotic and to reduce the systemic quantity, thereby decreasing general toxicity; they are whether industrially packaged or prepared in the operating theatre according to the antibiogramme [20]. The state of the art on how the antibiotic manages to act is the following: the antibiotic, when soluble in water, dissolves from the surface of PMMA into the tissues; antibiotic molecules of notable size are physically blocked inside the bone cement and, therefore, cannot spread from inside the cement to the surface. The dissolution process depends on the type of antibiotic, on the characteristic of the surface of the cement and on the way the cement itself is prepared. When the antibiotic is added to the cement during preparation of the cement itself, that is in the operating room, only a small part of the antibiotic molecules are casually on the surface of the cement and will be able to dissolute. This process explains why the actual antibiotic-loaded cements have a limited antiseptical action.