**4. Orthopaedic polymers**

### **4.1 Polyethylene**

One of Charnley's legacy to modern arthroplasty is the introduction of polyethylene as a bearing material for THA [14]. Polyethylene is a linear homopolymer,

composed of hydrogen and carbon, represented by the formula ~(CH2-CH2)~. Charnley originally used polytetrafluorethylene, termed Fluron G1 and Fluron G2 manufactured by Imperial Chemical Industries (London, England) as a bearing material for THA due to its low coefficient of friction and biocompatibility [23]. However, poor abrasive characteristics lead to failure of Charnley's polytetrafluorethylene acetabular cups within 2 years, due to low resistance to creep deformation. Fortuitously, Charnley's technician, Craven tested a material termed highmolecular- weight polyethylene (HMWPE) which was given to him by a plastic gear salesman. This material was first implanted in 1962 by Charnley when a HMWPE cup was used in combination with acrylic bone cement which was compressed into a reamed acetabulum (**Figure 3**) [24].

Ultra HMWPE (UHMWPE) is composed of long chains of polymerised alkene, ethylene. It is a semicrystalline polymer, having both a crystalline and amorphous phase. UHMWPE contains a set of ordered crystalline lamellae, with tightly packed randomly orientated macromolecules, embedded in a disordered amorphous phase. UHMWPE consists of up to 200,000 monomers per molecule and a molecular weight ranging from 2 to 6 106 g/mol. Increased molecular weight and degree of cross-bond

**Figure 3.** *Charnley hip consisting of HMWPE cup with a Moore stem [25].*

#### *Biomaterials in Total Joint Arthroplasty DOI: http://dx.doi.org/10.5772/intechopen.107509*

formation between chains increases the strength and wear resistance of UHMWPE. These properties combined with a low coefficient of friction, resistance to abrasive wear and corrosion, with a high impact strength and toughness have made UHMWPE the arthroplasty bearing material of choice. Comparatively high-density polyethylene has a lower molecular weight (0.05–0.25 106 g/mol) and higher density which results in inferior mechanical properties.

#### *4.1.1 UHMWPE processing*

Historically there have been 3 modes of UHMWPE production for orthopaedic implants. Direct moulding involves placing polyethylene powder into a mould of the final configuration of the implant. The powder is then placed under pressure and heated to consolidate it into its final shape. Ram extrusion involves a similar process; however, the powder is fashioned into a cylindrical bar stock which is later machined into the desired final shape of the implant. The final method involves moulding large sheets of polyethylene in which implants are later machined from. A more modern technique has recently been pioneered by Zimmer Biomet termed Hot Isostatic pressing. This process uses high temperature and pressure with argon gas and compression moulding. The process is then completed by a machining operation. These fabrication methods do not significantly alter the physical, chemical or structural properties of the original polymer other than the crystallinity. As the original polyethylene is heated to above melting point, its crystallinity is irreversibly decreased, decreasing fatigue strength. As a result, all polyethylene components possess their original properties prior to sterilisation.

#### *4.1.2 Sterilisation*

Polyethylene sterilisation presents a challenge as it cannot be carried out using traditional heating methods due it's low melting point. The main sterilisation methods used today include high energy radiation (gamma irradiation or electron beam irradiation) or surface treatment. Gamma irradiation is emitted from the decay of a 60Co unstable nucleus whereas electron beam is produced by heating a tungsten filament. Both radiation methods have the advantage of sterilising PE but also causing cross-linking of polymer chains which enhances fatigue strength and wear resistance [26]. However, electron beam irradiation can be performed in a shorter period of time (seconds) and with lower doses of radiation required to achieve a similar degree of crosslinking.

The main disadvantage of high energy radiation is oxidative degradation of the implant through radiolytic bond scission and free- radical generation [27]. Irradiation results in scission of chemical bonds of UHMWPE resulting in free radical formation. Environmental oxygen binds with free radicals permanently breaking the carbon-carbon bonds, a process termed oxidative degradation. This process results in reduced molecular weight and a final component with inferior wear resistance and increased wear debris generation. Macroscopically, oxidative degradation can be seen in retrieved and new components as appearing as a white band or crown effect, which results in severe failure, through delamination and fracture.

Several methods have been employed to reduce the effect of oxidative degradation during irradiation. Manufactures have trialled irradiating PE components in an inert gas such as argon or in a vacuum. Similarly, more sophisticated packaging methods have also become available with particle barrier material, preventing

ambient oxygen exposure. Packaging has also been designed to allow the penetration of select gases when chemically sterilised. During the 1990's, post-irradiation thermal stabilisation melting was also introduced as a means to reduce oxidative degradation by quenching residual free-radicals, allowing them to recombine, improving oxidative resistance. This was initially demonstrated to maintain mechanical property performance standards, but remelting decreased crystallinity, reducing PE fatigue strength.

Subsequently in 2007, UHMWPE stabilisation with antioxidant Vitamin E-diffusion was introduced in the United States [28]. Vitamin E diffusion, limited the need for post-irradiation melting, maintaining PE crystallinity. Synthetic α- tocopherol, the vitamin E used during implant processing, decreases the macro alkyl radicals available to interact with oxygen, inhibiting the ensuing oxidative cascade. Vitamin E can be incorporated into UHMWPE either by post-irradiation vitamin E diffusion or by mixing vitamin E with UHMWPE powder prior to pressurisation. Post-irradiation vitamin E diffusion has the benefit of avoiding the effect of Vitamin E on reducing the number of alkyl radicals available for cross linking but places the cross-linked PE at increased risk of oxidation prior to Vitamin E stabilisation. Mixing Vitamin E with UHMWPE powder has the converse effect.

## **4.2 Polymethylmethacrylate**

The earliest use of acrylic bone cement was by Glück in 1891, for use with an ivory hip prosthesis [6]. However, in 1901 a German chemist, Otto Röhm developed polymethylmethacrylate (PMMA), which is the earliest form of the bone cement which is in widespread use today in orthopaedics [29]. The Judet Brothers developed acrylic femoral hemiarthroplasties prosthesis as a treatment for hip arthritis [30]. However, it was Haboush who was the first to use PMMA as a grout to fix implant to bone [31]. Subsequently, Charnley pioneered to modern use of self-curing PMMA to achieve an implant- bone- cement construct for femoral and acetabular components, in the 1950s and 1960s. Charnley proposed that PMMA, creates an interface between prosthesis and bone to allow for distribution of contact forces and rigid fixation [14, 32]. PMMA acts as a grout which interdigitates with cancellous bone, enhancing interface shear strength. Cement may also act as a shock absorbing layer between elastic bone and a stiff implant with a Young's modulus (2400 MPa in tension) between cortical and cancellous bone. Cement therefore acts as an elastic interlayer between 2 stiff layers facilitating a more gradual transfer of stress from implant to bone. Currently, cement has a number of orthopaedic applications including; prosthesis fixation, percutaneous vertebroplasty, antibiotic delivery and as an interpositional material for bone defects.

PMMA (C5H8O2) is packaged as 2 separate components; a powdered polymer and a liquid monomer in a 2:1 ratio [33]. The liquid component is supplied as 20mls of fluid, in a brown vial to avoid the deleterious effects of direct sunlight. The liquid monomer contains methyl methacrylate monomer, consisting of microspheres of variable diameter with the size of particles determining the viscosity of the cement. Additionally, heat stable antibiotics may also be added such as gentamicin or vancomycin. Other compounds contained in the liquid component include N,N dimethyl- p toludine (DMPT), hydroquinone, and a colouring agent (e.g green chlorophyl or cobalt blue). The powdered polymers is typically packaged as 40 g of

powder containing pre-synthesised PMMA, barium sulphate or zirconium oxide and benzoyl peroxide.

Addition polymerisation of PMMA occurs via an exothermic reaction when the liquid and powdered components are mixed together. Benzoyl peroxide, in the powder component, acts as a catalyst, initiating polymerisation by interacting with DMPT in the liquid component [33]. This interaction liberates free radicals, breaking carbon-carbon bonds within MMA, allowing MMA monomers to couple with the growing polymer chain. Hydroquinone stabilises this reaction, preventing premature polymerisation. The radio-opacifier (barium sulphate or zirconium oxide) and colouring agent (cobalt blue or green chlorophyl) allow for identification of cement, radiologically and intra-operatively.

This chemical reaction is characterised by 4 distinct phases. The mixing phase or sandy stage (phase 1) occurs when powder and liquid components are mixed together (lasting roughly 30s). The waiting phase or sticky phase (phase 2), lasts approximately 1 to 3 minutes (depending on cement viscosity) and ends when the cement will easily separate from a gloved finger. The working phase (phase 3) occurs when the cement can be easily handled and signals when implants can be inserted, lasting approximately 5 minutes for high viscosity cement. The setting or hardening phase (phase 4) lasts approximately 1 minute 30 seconds to 2 minutes for high viscosity cements. These phases are influenced by a number of endogenous and exogenous factors. Increasing environmental temperatures and humidity decreases cement working time. Other factors such as a decreased powder liquid ratio increases setting time. The final biomechanical performance of cement can also be influenced by a number of exogenous factors such as mixing technique, sterilisation methods, antibiotics additives and choice of radio-opacifier.
