Silicone

Swanson marked the start of modern era of small joint arthroplasty with the development of the silicon spacer in 1966. ^{1 , 25 26 , 27} Silicones also known as polysiloxanes are polymers composed of repeating siloxane units forming chains of alternating silicon and oxygen atoms (Fig. 2.1). Silicone elastomers are created by three-dimensional (3D) cross-linking of linear silicone polymer chains. Chemical inertness, heat stability, and durability have led to its widespread use in medicine. Good biotolerance, high hydrophobic capacity, and ability to withstand various sterilization process are further benefits allowing use of silicone elastomers in implant arthroplasty. ^{26 , 28 29 30 31 32 , 33}

Hardening of silicone elastomers is by curing, a chemical process which adds curatives to induce polymer cross-linking. ²⁹ Tensile strength is further enhanced by incorporation of “filler.” Silicone elastomers for medical applications normally utilize amorphous silica filler to reinforce the cross-linked matrix. While mixing the silica with the silicone polymers, a hydrogen bond is formed between hydroxyl groups on the filler’s surface and silicone polymer, resulting in higher tensile strength and better capacity to resist elongation (Fig. 2.1). ³⁴

The original Swanson prosthesis was made from heat-vulcanized, medical-grade silicone elastomer stock. The prostheses are single-piece and formed in a mold at 121.1°C under 22.8 kg/cm² of pressure. ¹ Most silicone finger prostheses consist of intramedullary stems bridged by a hinge ²⁶ (Fig. 2.2). The inherent flexibility of the silicone elastomer allows flexion and extension at the hinge and provides dampening effect on the bone. At the same time, it is stiff enough to maintain some joint alignment. The long-term functional stability of silicone prostheses occurs by development of well-encapsulated fibrous capsule which starts to develop within 3 days of implantation. ³²

However, there are problems with the silicone elastomer implant. These include implant fracture, wear debris generation, and particulate synovitis.

Joyce et al analyzed the cause of silicone implant fracture by retrieving 12 Sutter metacarpophalangeal (MCP)prostheses from three patients. ³⁵ They found that fractures generally occurred at the junction of the distal stem and hinge as a result of subluxing forces concentrated on MCP joint and flexion occurring at the stem rather than the hinge. In addition, the cortical bone of the proximal phalanx impinges on the dorsal surface of the distal stem of prosthesis. It is presumed that sharp spurs from the proximal phalangeal bone will produce small cuts over dorsal surface of the prostheses. With time, the crack propagates resulting in a fatigue fracture (Fig. 2.3). ^{35 , 36} Two-thirds of silicone MCP joint prostheses fracture by 14 to 17 years after implantation. ^{37 , 38} Fortunately, by the time a fracture occurs, a fibrous pseudojoint is formed. The finger joint has some stability and often do not require revision surgery. ²⁶

During the process of abrasion and fatigue fracture, microscopic wear particles shed around the joint. This will stimulate proliferation and accumulation of mononuclear cells which will secrete cytokines and proteolytic enzymes, causing swelling and inflammation of pseudosynovial and synovial membranes. Persistent chronic synovitis can lead to fibrosis, osteolysis, and bone necrosis surrounding the prosthesis. Despite these downsides, silicone arthroplasty remains popular options in finger joint arthroplasty.

Ultrahigh Molecular Weight Polyethylene (UHMWPE)

Polyethylene (PE) has the advantage of easily being formed into many different shapes. It is mainly used coupled with a metallic component. It is formed from ethylene gas and is polymerized by Ziegler–Natta catalyst (titanium(III) chloride) into UHMWPE powder. The powder is consolidated under high temperatures and pressure. Because of its high melt viscosity, the final product is produced by compression molding and ram extrusion. ⁴¹

UHMWPE is a high-density PE having a semicrystalline structure with a molecular mass more than 2,000,000 amu. The macromolecules consist of local ordered sheetlike crystalline lamellae embedded within amorphous regions and communicate with surrounding lamellae by tie molecules (Fig. 2.5). ⁴² It has been extensively used in large joint arthroplasty because of its low friction, high resistance to wear, and high toughness. ⁴³

Medical-grade UHMWPE is free of calcium stearate and requirements are set according to the American Society for Testing and Materials (ASTM) 648–14. ⁴⁴ Since the polymerization of UHMWPE requires a specialized production plant to handle the dangerous chemicals, there are only two companies capable of making these resins (Table 2.3). Despite identical manufacturing methods, there are slight variations in molecular weight and resin morphology which affects their mechanical properties and wear abrasion resistance. ⁴⁵

The main factor responsible for UHMWPE failure in joint replacement is oxidative degradation. The presence of oxidation is related to the sterilization and storage methods. High-energy radiation such as gamma or electron beam radiation are the most commonly used sterilization methods for PE components. ⁴⁶ A nominal dose of 25 to 40 kGy gamma radiation is commonly used, which will emit energy higher than that of the polymeric chemical bonds. This would generate the scission of some chemical bonds of the UHMWPE, and decrease its molecular mass. In the presence of oxygen, this will lead to the formation of free radicals, causing worsening of some chemical and physical material characteristics. ^{47 , 48 , 49} The oxidation process continues as long as there is an oxygen supply. This phenomenon is known as post-irradiation aging. It has been shown to occur in UHMWPE implants that were gamma sterilized in air and packaged in air-permeable packaging. ^{50 , 51 , 52} Decreased abrasive wear resistance, due to oxidation, leads to a decrease in mechanical properties such as abrasive wear resistance which results in the formation of wear debris and consequently osteolysis, which has been recognized as being the main cause of failure in orthopaedic implants. ⁵³

To decrease the effect of oxidation on wear and the mechanical properties of UHMWPE, orthopaedic implant manufacturers modified their sterilization protocols aiming to reduce the amount of oxygen exposure during storage. These include gamma radiation sterilization in vacuum-packaging or inert-gas packaging. However, free radicals that are already present cannot be eliminated and in vivo oxidation is still possible. ⁵⁴ Therefore, several methods have been attempted to improve the wear resistance of UHMWPE, including cross-linking, thermal treatment, and the addition of antioxidant vitamin E.

Cross-linking is achieved by formation of active sites at the chain ends, which can recombine to form trans-vinylene bonds. ⁵⁵ This can be achieved by gamma radiation and the peroxide method. The wear rate of highly cross-linked UHMWPE is six times less than conventional UHMWPE. ⁵⁶ These should allow manufacturers to produce thinner inserts that could be applied in hand arthroplasty. However, as the cross-linking density increases, elastic modulus, ultimate tensile strength, yield strength, and ductility reduce, which reflect the reason why cross-linked UHMWPE has less resistance to fatigue crack propagation. ⁵⁶ Therefore, it is preferable to have optimal cross-linking density and keep it local to the surface to retain the bulk mechanical properties. ⁵⁷

To remove free radicals formed during the cross-linking or sterilization of UHMWPE, thermal treatment has been employed. Two common thermal treatments are remelting and annealing. Remelting is when the temperature is elevated above the melting point (150°C), which allows free radicals trapped inside the crystalline region to diffuse out. Typically, the residual free radicals drop to undetectable level with this method. ⁵⁴ But remelting has been shown to significantly reduce the degree of crystallinity, which affect the mechanical properties. ^{58 , 59} During annealing the temperature does not exceed the melting point, so the crystalline region does not undergo dissolution but this leaves measurable amounts of free radicals inside processed UHMWPE, which is still susceptive to in vivo oxidative degradation. ^{60 , 61}

Vitamin E is a natural antioxidant, which is able to consume the free radicals and improve wear resistance. Two methods of incorporation are currently used: one is to blend vitamin E (concentration less than 0.3 wt.-%) with UHMWPE powder before consolidation; and the other is to allow vitamin E to diffuse into the bulk UHMWPE. Vitamin E–containing UHMWPE has shown superior oxidative stability with greater mechanical and fatigue strength; these are now used clinically. ^{62 , 63 , 64}

Polymethylmethacrylate (PMMA)

Polymethylmethacrylate (PMMA) is commonly known as bone cement. It has been the material of choice for implant fixation to the host bone over 50 years, after being popularized by John Charnley in 1958. ⁶⁵ The primary function of PMMA is transfer of weight-bearing forces from bonetoimplant and vice versa. The term “cement’ is a misnomer because PMMA has no adhesive properties. Rather, it acts as a space-filler to create a close mechanical interlock between the rough medullary canal and the implant. This close association leads to optimal stress and interface strain energy distribution which are paramount in longevity of the implant. If the extreme stresses generated exceed the capability of the bone cement to transfer and absorb forces, fatigue fracture will result.

PMMA is an acrylic polymer that is formed by mixing liquid methyl methacrylate (MMA) monomer and a powdered MMA-styrene co-polymer. Benzoyl peroxide (BPO) in polymer powder acts as initiator while N,N-dimethyl para-toluidine (DMPT) or dimethyl para-toluidine (DMpt) in liquid monomer act as an accelerator for the free-radical polymerization reaction at room temperature (cold curing). To avoid premature polymerization of the liquid component during storage, hydroquinone is added as stabilizer (Table 2.4).

Most of the commercial bone cements have a polymer powder/liquid monomer ratio of 2:1 (weight in g/volume in mL) in order to decrease the shrinkage level due to the density change in MMA-to-PMMA conversion during the polymerization. When mixed together, the reaction starts to take place between BPO initiator and the DMPT accelerator forming benzoyl free radicals at room temperature. These unstable free radicals start the initiation step of polymerization by forming a new chemical bond between the initiator fragment and one of the double carbon bonds of the monomer molecule. A new free radical is created on the activated monomer molecule which can react with the double carbon bond of a new monomer molecule in the same way as the initiator fragment did. This process, known as propagation, takes place over and over again to form a long chain of monomeric units. Polymerization ceases when there is regrouping of the two radical chains leading to depletion of free radicals (Fig. 2.6). ⁶⁶

The free-radical polymerization process of PMMA bone cement is a highly exothermic chemical reaction as a result of energy released from the breaking down of C=C double bonds during polymer chain propagation. The invitro peak temperature can reach 110°C. It is particularly influenced by the liquid monomer composition, the polymer powder/liquid monomer ratio, and the radio-opacifier content. ⁶⁷ Bone necrosis can occur when exposed to temperature of 50°C for 1 minute or 47°C for 5 minutes. ⁶⁸ Although clinical studies have showed lower peak temperatures (40–47°C) has detected at the bone interface, which may be due to a thinner cement coating and heat dissipation through the broad implant surface and local blood circulation. ^{69 , 70} However, thermal-induced bone necrosis has been shown in animal model study and the possibility of leading to early loosening and subsequent implant failure are still the concern. ⁷¹

Various additives improve the workability of the polymer (Table 2.4). To make the cement radiopaque, contrast agents such as barium sulfate (BaSO₄) or zirconia dioxide (ZrO₂) are added. Chlorophyll dye helps to distinguish bone cement from native bone and facilitate removal when performed revision surgery. Antibiotic-loaded bone cements are often used as drug-delivery systems. Implants are more susceptible to bacterial colonization on their surfaces because the bacteria can form a glycocalyx-blocking immune scrutiny by the body and leading to periprosthetic infection. The local active antibiotic levels are often not sufficient to reach a therapeutic dose when delivered by the intravenous route alone. When loaded with antibiotics, bone cement functions as carrier matrix and offers a synergistic effect with systemic antibiotics. ⁷² Many antibiotics has been added to PMMA bone cement clinically but many factors including bacteriologic, physical, and chemical aspects need be assessed to optimize outcomes (Table 2.5). ⁷³

During different stages of PMMA bone cement polymerization, the appearance and handling characteristics change. There are four different phases. The mixing phase is the period when thorough mixing of the powder and liquid component takes place. It can be mixed manually or with a commercially available mixing system with the ability to apply a vacuum to minimize the number of pores; this phase lasts around 1 minute. In the next (waiting) phase the viscosity of bone cement progressively increases to a condition that can be handled without sticking to surgical gloves; it lasts for several minutes. The bone cement surface often shows the distinctive change from a gloss to a matt appearance. The working phase is the period during which the surgeon can apply the cement and insert the prosthesis; this lasts around 2 to 4 minutes. The viscosity should be high enough to withstand the bleeding pressure but fluid enough to allow cement interdigitation into the surrounding cancellous bone. These two phases often depend on the type of cement and the handling temperature. Bone cement hardens when reaching the setting phase; this is the exothermic period with peak temperatures. The overall setting time, starts from the time of mixing liquid and powder to the time when the cement temperature drops to halfway between the ambient and maximum temperature; it usually lasts 8 to 10 minutes. Many factors can affect the setting time (Table 2.6). ^{74 , 75 , 76}

Some of the designs of MCP and PIP joint surface replacement require cement to help to accommodate the variations in endosteal canal configuration and to restore proper alignment. ⁴⁰ PMMA bone cement serves as a suitable material to fill up the defect between the implant and bone because of its excellent intraoperative workability. However, hand joint arthroplasty systems do not have any cement restrictor or cement gun to achieve pressurization, meaning there will be a less uniform and so weaker cement mantle that is prone to developing overstress. This may lead to cement fracture with release of cement particles which would induce a local inflammatory reaction and possibly periprosthetic osteolysis. ⁷⁷

Metals Alloys

Metals alloys are constituted from metallic and nonmetallic elements which are frequently used because of their mechanical strength. Most of the metals used for biomedical applications comply with the ISO 5832 standard. They mainly include three families of metals: stainless steel; cobalt chromium (CoCr) alloys; and titanium (Ti) alloys. ¹⁹ They have different mechanical properties and corrosion responses. All metal alloys are stiff, ductile, and hard. Appropriate metal selection for a specific application requires consideration of a variety of parameters.

Pure metal forms a crystalline lattice microstructure, i.e., a repeating 3D unit with a closely packed, high-density arrangement of metal atoms. They are packed into one of three crystalline arrangements: body-centered cubic, face-centered cubic, or hexagonal close packed (Fig. 2.7). However, the units often do not combine perfectly and form clusters of grains, imprecisely aligned with each other. This causes imperfections within grain microstructure and defects in macrostructure known as scratches and voids. The size of the grains affects the mechanical properties of strength and isotropy on the materials. Biomechanically the larger the grain size, the earlier material fatigue and failure. ¹⁴

Alloys are usually formed by adding the alloying metal elements to the principalmetal element in the molten state. When it solidifies, the metallic alloying elements, such as chromium and nickel in stainless steel, substitute for iron atoms in the crystalline lattice because of similar atomic sizes. When nonmetallic alloying elements are added, they are smaller and so fit in the spaces among the metal atoms. Deformation of the crystalline lattice occurs, which, in turn, causes resistance to the flow between crystal units leading to increase in resistance to plastic deformation and yield stress. ¹⁴

Several steps are required to turn raw metal alloys into metallic components for joint prostheses. Casting is regularly used in manufacturing orthopaedic implants as it provides an efficient way for mass production that is difficult to achieve by machining. But the casting process produces large grains and metallurgical imperfections so they have lower mechanical properties than wrought or forged alloys. Forging is a manufacturing process to optimize material properties under pressure and classified as cold or hot working relative to its recrystallization temperature. Cold working increases the number and density of strains by plastic deformation at room temperature which makes it difficult to deform further. The yield strength increases but ductility decreases. Hot working is plastic deformation above the recrystallization temperature; it allows an increase in grain size which increases homogeneity and ductility. Sintering is the process of compacting and forming a solid mass of material by heat or pressure without reaching the melting point of the material. It can increase the mechanical strength of the material and is used for the shaping process of materials that have extremely high temperature. A summary of different techniques producing the metallic parts of joint prostheses from various alloys are listed in Table 2.7. ^{14 , 78 79 80}

Calcium hydroxyapatite (HA)

Calcium hydroxyapatite is a ceramic. Ceramics are compounds of metallic elements bound ionically and covalently with nonmetallic elements. The calcium-to-phosphate ratio of HA resembles the minimal phase of bone, which accounts for their osteoconductive features. ⁹⁸ Once HA is implanted in vivo, a layer of carbonate-hydroxyapatite (CO-HA) will develop on its surface as a result of ion exchange with the surroundings. ^{99 , 100 , 101 , 102} The resulting coating layer on the HA is known as carbonate-hydroxyapatite (CO-HA); it resembles the apatite present in normal bone. ¹⁰³ The new apatite layer acts as a scaffold for osteoblasts and with time, it is further resorbed by osteoclasts and replaced by new bone. ¹⁰¹ Some studies have shown that HA-coated implants reduce the migration of components and have higher survival rate of comparable press-fit implants. ¹⁰⁴ However, because of its low interface bonding strength and low toughness of the HA layer, it can result in fracture and delamination from the stem surface leading to osteolysis or third-body wear if the debris migrate into the joint space. ¹⁰⁵ A different coating technique has been investigated with the aim to improve the interface bonding strength of the HA layer (Table 2.8). ⁷⁹