Stainless “surgical” steel remains one of the most frequent applied alloys to manufacture surgical implants and instruments. These alloys serve for the fabrication of at least half of all orthopaedic implants used in the USA [8], although they are gradually being displaced from the market by other alloys, notably CoCrMo and titanium alloys.

Stainless steel possesses several desirable properties. This material is durable, ductile, and, thus, relatively simple to process. It is also non-toxic and biocompatible, as it does not evoke an adverse reaction from adjacent tissues. Technology of its production (smelting, casting, and forging) and processing (cold-hammering, tempering, machining, and threading) is well known and relatively low cost. The final products made of steel are available in practically unlimited quantities for an acceptable price.

Nowadays, an austenitic 316 L steel is mainly used for implants due to its high corrosion resistance. It consists of reduced carbon (below 0.03%) and increased chromium (16–18%) and nickel (10–14%) content, with the addition of molybdenum (2–3%), manganese (ca 2%), and small additives of sulphur, silicon, phosphorus, and nitrogen [9]. Corrosion resistance results from a thin Cr₂O₃ layer, that passivates on the outer layer. Such layers protect the human organism as the implant does not interfere with metabolic processes, that occur in the body. It should be mentioned, however, that there is the possibility for exceedingly high chromium and nickel content to lead to several unwanted side effects, as both may irritate tissues and lead to immune reactions. This has been demonstrated in the literature where up to 20% of the population of industrialised countries demonstrate sensitivity to the chromium and nickel [9]. Chromium and nickel may be carcinogenic, and their high concentration may even be toxic [10,11] and may promote infections caused by nickel-dependent bacteria. Nickel itself promotes ingrowth of those organisms being a part of several microbial metalloenzymes [12,13].

Other steels, including 200, 400, and 500 series, are also used for biomedical applications due to their reduced chromium concentration (especially 500 series steel), although less frequently than the 316 L steel.

The mechanical properties of 316 L steel predispose it for various medical applications including: pins, rods, intramedullary nails, screws and plates, and even joint prostheses. Although 316 L steel is highly resistant to corrosion, it is susceptible to stress cracking and crevice corrosion. The first originates from access to chlorides that have been found in biological fluids. The second, from the fracture of the ultrathin, protective oxide layer that passivates the material outer surface. The susceptibility to stress corrosion cracking increases with exposure to chloride-rich, biological environments. Such exposure requires stainless steel implant removal as soon as they fulfilled their function, thus reducing the material to trauma procedure applications. In addition, crevice corrosion may occur when the implant succumbs to intermittent bending. The fracture of the oxide enables the corrosion of core material resulting in the deterioration of its mechanical properties and subsequent failure. In order to protect the implant material against crevice corrosion, two approaches were recommended: implementation of a thicker oxide passivation layer on the surface, and a careful application to avoid oxide fracture. The first approach relies on special preparation of the implant’s surface. Polishing smoothens its surface, thus reducing the contact with the outer environment; chemical preparation with nitric acid thickens the oxide layer. On the other hand, electrolytic passivation (anodizing) removes free iron particles from its surface, locally increasing the concentration of chromium and nickel, that are responsible for the resistance against corrosion [14]. The second approach requires an appropriate technology of the implant manufacturing and handling. Casting or forging form the final product into the desired shape with specific mechanical properties. Casting enables complex shapes to be produced and is a relatively simple and low-cost process. On the other hand, labour-intensive and costly forging allows for the production of an implant that is much more ductile and durable. It should be mentioned, however, that cold-working strengthens the material, but also increases brittleness. Hence, the cold-worked implants (i.e., intramedullary rods), used to stabilise shafts of a long bone, are more durable, but are not suitable for bending loads as much as cast implants.

In summary, to obtain a defect free passivation layer covering an implant, several conditions should be maintained. First of all, implants should be properly designed to guarantee an appropriate stiffness, in order to withstand mechanical loading occurring in typical loading scenarios. Surgeons should perform stabilisations without tampering with the structure; that is, without the need for bending to adjust to the bone shape. It is noteworthy, that each bend of the implant, especially cyclic one, disrupts or deteriorates the passivation layer properties and leads to implant fracture. Additionally, special stabilisation techniques (tension band principle) defined by Pauwels in the 1930s is commonly used to reduce the risk and amplitude of plate bending during limb loads, based on the conversion of tensile forces acting over the fracture compared to compressive loads [15]. Thus, both the production and implantation technique provide the desired mechanical properties and protection against crevice corrosion that protect from unwanted electrochemical processes, resulting in loss of durability, and subsequent fracture.

Titanium and its alloys have been known since the end of the 18th century. Pure titanium, used to fabricate several alloys characterised by a relatively high hardness and corrosion resistance, found its first medical application in the 1940s as dental implants. These alloys were also used in orthopaedics due to desirable mechanical capabilities and the ability for osseointegration, defined as the capacity to bind with adjacent bone, improve implant stability and reduce the risk of losing the implant [16,17]. Additionally, a high corrosion resistance enabled the implant’s adoption for several decades without any obvious tissue irritation or toxicity effects [18]. With increased demand from aircraft factories and submarine shipyards, global production for titanium rapidly increased in the 1950s and 1960s, enabling expanded applications in the medical field. The first publication that discussed the possibility of using titanium as a surgical implant dates to 1963 [19]. The study brought an increased interest on the subject in the subsequent decade [20,21]. Low density, high strength, and high corrosion resistance predispose this metal to the production of surgical implants, especially in its beta allotropic form, and alloyed with molybdenum, vanadium, niobium, tantalum, and zirconium. Nowadays, titanium alloys have been widely used to fabricate trauma and orthopaedic implants [22,23] due to increased biocompatibility, lack of toxicity, osseointegration, high tensile strength to density ratio, and corrosion resistance. The most popular titanium alloy used for implants is aluminium-vanadium doped alloy (Ti6Al4V). Currently, practically every type of orthopaedic implant has a titanium ‘variant’, including screws, plates, intramedullary nails and rods, external fixators, and joint prostheses. Since titanium is non dielectric and does not increase in temperature when exposed to alternating magnetic fields, it is ideal as an implant as it also does not interfere with magnetic resonance imaging [24]. This significant advantage of titanium has dominated the materials’ application in traumatology and joint replacements, and practically monopolised the market of implants used in spine surgery [25,26]. Additionally, its elasticity is much more comparable to the viable bone rather than that of the steel. The similar properties of implant and bone enable to avoid non-desired strain components and an overload at the bone-implant interface, thus reducing the risk of loss or periprosthetic fracture [27].

To date, there is minimal evidence to suggest immune adverse reactions from titanium implants, although the possibility to activate discrete cellular reactions has been postulated, as activation of leukocyte emigration and their concentration at tissues adjacent to titanium implants have been observed in the literature [28]. An interesting finding is that leukocyte emigrations were not as severe around stainless-steel implants, possibly due to their high nickel content [29].

Titanium is rarely used in its pure form [30]. Nevertheless, it still serves for an implant’s coating with spongy, three-dimensional, plasma-sprayed layers, that provide at least some titanium characteristics to other materials [31]. In the vast majority of cases, the Ti6Al4V and its derivatives are used in orthopaedics. However, newly designed alloys, including TiNbZrTaSiFe [32], TiMoFe [33], and TiMoNbZr [34], are characterised by the modified or improved mechanical properties and have become an alternative to traditionally used alloys. The new generation titanium alloys exhibit greater elasticity (e.g., the Young’s modulus ca 50–65 GPa) that is similar to that of bone, which predispose them as a more suitable material for orthopaedic purposes. To manufacture intricate components from these new alloys, novel methods have increasingly been studied including the methodology of personalised, computer-designed, 3D implant “printing” using laser-beam sintering technology [35]; however, the enormous potential of this method has not been widely adopted to a large scale.

316 L austenitic steel has been found to be susceptible to wear due to friction between working parts of an implant. Hence, wear resistant materials, including CoCr alloys, have been applied, often produced with some content of Mo and other metals including nickel, tungsten, and titanium. Specifically, the most common orthopaedic implant alloys contain between 62–68% Co, 27–30% Cr, 5–7% Mo, and <2.5% nickel, with an example alloy classification used for medical purposes being ASTM F75 CoCr alloy [36,37,38].

CoCr alloys were introduced in early 1900’s and have been characterised by good biocompatibility, high wear, and corrosion resistance, which result from high cobalt, molybdenum, and chromium content (almost twice that of steel). Moreover, these materials are simple to cast, and, thus, complex shaped implants could be produced at relatively low cost, without requiring further surface treatments compared to stainless steel. Thus far, several implants and medical instruments have been manufactured from CoCrMo alloys, including surgical blades and needles, cardiac valves, cases of pacemakers, and joint and dental prostheses. The material has exhibited excellent performance for working parts of joint implants, including heads of hip and condylar components for knee prostheses.

Vitallium, introduced in 1939, is one of the most popular CoCrMo alloys (65%, 30%, and 5% wt., respectively) used for the manufacture of joint replacements, starting from Charnley’s hip prosthesis [36]. It was found to be extremely durable, with orthopaedic implants manufactured from this material being in continuous use for as long as 70 years [37]. Unfortunately, the implants are susceptible to breaking during bending, showing their limited usefulness in long bone fracture stabilisations. Another disadvantage is the relatively high chromium content jeopardising immune reactions, as the percentage of the population sensitive to this metal in modern societies has increased. Nevertheless, high wear resistance, good biocompatibility, and low cost of manufacture have made CoCr alloys very popular for orthopaedic implants in the 1960s [38], with subsequent loss of interest resulting in its replacement by titanium alloys, when the number of adverse effects was found to increase [39,40]. This accelerated, when the toxicity of wear debris produced by metal-on-metal prosthesis became well known [41].

Aluminium and zirconium oxides (Al₂O₃, ZrO₂) and mixed oxide ceramics are used to manufacture working parts of joint prostheses components. CoCr alloys are characterised by high stiffness, scratch and corrosion resistance, and good biocompatibility. The technology of their production is relatively simple and low cost. Thus, several manufacturers offer implants of/for ceramic-on-ceramic articulation systems.

The ceramic acetabulum and prosthetic head ensure low friction and a small amount of wear debris. However, they are exposed to fragmentation, when succumbing to mechanical overload. The ceramic joints could also produce an irritating squeaking while walking [60,61].

Primarily, alumina ceramic was the most extensively used, being replaced by zirconia due to its higher endurance and lower susceptibility to fracture. It should be highlighted, however, that all ceramics are predisposed to brittle failure when subjected to excessive mechanical loads. Thus, polyethylene inserts were introduced to reduce those loads. In a configuration with ultra-high-molecular-weight polyethylene (UHMWPE) acetabular insert ceramic head of the hip prosthesis exhibits reduced risk of fragmentation [61].

Recently, mixed alumina (Al₂O₃) and zirconia (ZrO₂) ceramics, and those stabilised with yttrium oxide (Y₂O₃) or lithium silicate (Li₂SiO₃) were brought to the market. These ceramics are characterised by considerably higher toughness and fragmentation resistance [62,63]. Pure ZrO₂ was found to be very brittle during the production process and cooling in particular. Thus, manufacturers alloy the material with stabilisers (calcium, magnesium, yttrium, and cerium oxides; CaO, MgO, Y₂O₃, and CeO₂) which enable more durable yttria-partially stabilised tetragonal zirconia polycrystals (Y-TZP) to be obtained. Due to its biocompatibility and mechanical properties, it was found to be suitable for dental applications, although too fragile for orthopaedic implant manufacturing [64]. Thus, for orthopaedic purposes, Y-TZP is usually reinforced with Al₂O₃ forming alumina-toughened zirconia (AZT) that is much more resistant to cracking than Y-TZP [65].

Tetrafluoroethylene or polytetrafluoroethylene (PTFE), better known as Teflon or Syncolone, is a synthetic fluoropolymer (C₂F₄)_n that was first manufactured in 1938. Together with its expanded form (ePTFE; Gore-Tex), it found a wide range of applications due to several unique properties [69,70,71,72,73,74]. From an orthopaedic point of view the most important are the mechanical properties. The material is non-stick and highly slippery, thus, significantly reducing the friction between working parts. Moreover, it is well tolerated in between tissue due to its extreme non-reactivity, corrosion resistance and biocompatibility [69]. Medically, Teflon was primarily used to manufacture endovascular and urinary catheters, vascular, biliary, and ocular prostheses, and as a material for soft tissue reconstructions [70]. It was used in orthopaedics to reconstruct ligamentous [71] and tendinous defects [72]. Teflon is also used in arthroplasty [73] and even to stabilise bone fractures [74]. The latter application seems to be questionable due to the inappropriate strength of the material. However, as a lubricant that reduces the friction between working parts of an implant, it seems to be an unarguably excellent alternative to other fluorine-based plastics, including ethylene tetrafluoroethylene (ETFE; brand names: Fluon, Tefzen and Texlon) [75].

Polyethylene (PE; (C₂H₄)_n) is, nowadays, the most extensively used plastic in the world. It is a linear homopolymer consisting of hydrogen and carbon. It is a tough, abrasion and corrosion resistant, bioinert, self-lubricating, slippery, and semi crystalline polymer. It is also characterised by the density of 0.93 g/cm³, yield point of 20 MPa, and Young’s modulus of 700 GPa [76]. Polyethylene was first synthesised in 1898 by von Pechmann, while working on diazomethane [77]. Nevertheless, an attempt to synthesise it on an industrial scale was carried out in 1933 by Fawcett and Gibson. They polymerised free radicals under high temperature and pressure obtained the low-density polyethylene (LDPE). LDPE is still used for the production of plastic bags, packaging foams or plastic wraps. From the 1950s, the synthesis of the polyethylene proceeded under low pressure and temperature due to the elaboration of polymerisation catalysts. As a consequence, the high-density polyethylene (HDPE), characterised by increased hardness and tensile strength, but decreased elasticity when compared with LDPE, was obtained [78]. Neither forms were suitable for orthopaedic purposes due to their inappropriate physical properties. Thus, ultra-high molecular weight polyethylene (UHMWPE), with chains consisting of up to 200,000 monomers per molecule (HDPE only ca 1700) and molecular weight from 2 to 6 million g (HDPE: 0.05–0.25) was synthesised [79,80,81]. UHMWPE is predisposed to manufacture acetabular cups of hip and inserts of knee prostheses, as well as artificial, intervertebral discs due to good strength-to-weight ratio, low moisture absorption (almost none), extremely high impact strength, and resistance to abrasion from the high degree of polymerisation [82]. Moreover, UHMWPE is approximately 15 times more resistant to abrasion than steel and has a lower friction coefficient. Its production is also simple and cheap. It should be noted that desirable physical properties of UHMWPE, including resistance to tensile loads and shear stress, are associated with very long chains and their intermolecular attractions induced by Van der Waals forces [81].

UHMWPE cup with acrylic bone cement was firstly attached to the reamed space of the hip’s acetabulum by Charnley [83]. Throughout the history of UHMWPE, several attempts have been made to reinforce this polyethylene. In the 1970s, Zimmer developed polyethylene reinforced with carbon fibres. Unfortunately, it was characterised by inferior properties including reduced material strength and wear resistance in comparison to the original UHMWPE. In the late 1980s, an additional effort was made to reinforce UHMWPE by its high pressure recrystallisation (DePuy). However, the new material, called Hylamer, was of lower strength than the original UHMWPE. Its application was discontinued in the second half of the 90’s. It should be mentioned, however, that failures of Hylamer were found to be associated with radiation sterilisation [79].

The breakthrough in polyethylene production occurred in 1998, when crosslinked UHMWPE was synthesised. Its low friction, improved mobility, reduced wear debris, and greater elasticity than most metals significantly lowered the risk of a loose implant. Polyethylene is now used in various types of joint prostheses to manufacture components working with metal and ceramic materials. Despite the fact that UHMWPE performs well as a material for moving parts of endoprostheses, its abrasive products (wear debris) activate osteoclastic bone resorption stimulating the implant’s loosing [84]. It is also exposed to creep, resulting in the implant’s deformation [85], which requires further revision procedures.

Polimethylmetacrylate (PMMA; C₅H₈O₂) was primarily introduced into neurosurgery and dentistry in the 1940s. For orthopaedic purposes, Judet elaborated acrylic implant reinforced with a metallic pin to restore the femoral head [86]. Charnley used the self-curing PMMA as a bone cement to anchor the prosthetic stem made from metal (1960). Currently, this material is used in orthopaedics to fix components of joint prostheses, and in the surgical treatment of osteomyelitis and infectious complications of orthopaedic implants, as well as in vertebroplasty and kyphoplasty. PMMA is also commonly used to strengthen the anchorage of implants in osteoporotic bone and to reconstruct metastatic bone defects. It is characterised by relatively low density of 1.18 g/cm³, ultimate tensile strength of 72 MPa, Young’s modulus of 310 GPa, and elongation of 5% [87,88]. PMMA is hard, stiff, brittle, and possess limited adhesiveness. It is usually used as a grout filling in the narrowed spaces of the bone marrow cavity in osteoporotic bone or attaching the implant to the cancellous bone filling in free spaces of its pores. When it is compressed, especially by shock forces of high amplitude, or undergoes severe bending, PMMA may break [89].

Polymerisation of PMMA proceeds as a chemical reaction between two components at room temperature. The first component, initiator, usually methylmetacrylate (MMA) or polimethylmetacrylate (PMMA) used as an amorphous powder mixed with radiopaque (e.g., barium sulfate; barite), when mixed with the second component, usually a liquid activator (MMA monomers mixed with stabiliser), begins to polymerise, forming amorphous PMMA. The process is exothermic and usually takes several minutes, where the temperature may rise up to 82.5 °C. PMMA monomers (MMA) are highly irritant and even carcinogenic even though PMMA itself is biocompatible and does not evoke tissue irritations. These materials could also lead to hypotension and lung fat embolisation. Thus, a precise amount of initial component material must be used to minimise the risk of side effects produced by unbounded monomers.