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Biomaterials for hip implants – important considerations relating to the choice of materials

Biomaterials for hip implants – important considerations relating to the choice of materials IntroductionThe first hip joint replacement took place in 1926 in the USA. Hip joint replacement is a major achievement in orthopaedic surgery in the 20th century. Scientific research and experiments with a large number of human body part implants are carried out. Medicine, mechanical engineering, chemistry, biology, physics and material science have created particular divisions and branches intended to investigate implants.The synergy of medicine, biochemistry, biophysics, mechanical engineering, materials science and informatics succeeds in development of real progress in orthopaedic surgery of hip joint implants. Recently, several millions of hip joint replacements are made per year in the USA. Similar progress takes place in many other countries. Among others, the quality of the materials used and the material science, in general, are involved in achievements in this interdisciplinary field. Therefore, this scientific field with respect to hip joint implants is discussed and considered in this article.Most parts of hip joint implants are made of metallic materials due to heavy and cyclic load bearing. Additionally, they work in a bioactive environment. Titanium alloys are the most advanced materials in this type of application; however, Co-Cr-Mo alloys and austenitic stainless steels as bio-steels mated with appropriate metallic alloys, ceramics and polymers are the main materials of hip implant components. Femoral heads and sockets (cups) can also be made from advanced ceramic materials. These statements are based on experience and scientific biomedical investigations [1], [2], [3], [4], [5], [6], [7], [8], [9], [10]. Additionally, ultrahigh molecular weight polyethylene (UHMWPE), high-density cross linking polyethylene and polieteroeteroketon are applied to produce the circular bearing.Metallic alloys are crystalline materials since their properties strongly depend on the type of the crystal lattice. The specific space atomic arrangement defines the crystal lattice, which in turn relates to the particular properties and their anisotropy. The chemical and phase composition of an alloy as the main characteristic has a major influence on the crystal lattice, grain size, lattice defects and crystalline texture. The microstructure revealed with microscopic investigation provides information about the shape and size of grains and precipitations, which is the third important characteristic of biomaterials. All these properties of biomaterials are formed during technological operations of production like the metallurgical process, solidification, sintering, metal forming (rolling and forging), machining, heat treatment and coating deposition.All these technological processes are optimized in order to achieve the optimal structure and microstructure coupled with the expected mechanical, chemical and biological properties. The metallic components usually undergo additional surface treatments including coating and alloyed surface layers deposition.Recently, titanium alloys are most often used in biomaterial applications due to an excellent combination of properties. The titanium alloys are characterized by good mechanical properties, low density, an acceptable tissue tolerance, high strength to weight ratio, outstanding resistance to corrosion by body fluids, high biocompatibility, being nonmagnetic and the capacity for joining with bone, i.e. osteointegration [1], [4], [11]. Amongst others, it is important that titanium is a material that will not introduce fibrous tissue barriers in contact with a healthy bone. Another advantage of titanium is the fact that the fatigue properties of the load bearing device are not reduced through contact with body fluids containing aggressive chlorine ions. The biological, chemical and mechanical contacts of foreign material with the bone, body fluids and tissues are most important characteristics of the implant, although the interaction between mechanical parts of a complex and modular implant is an additional serious area of scientific investigations. The components of modular implants (Figure 1) are exposed to degradation processes as fatigue loading, contact fatigue, corrosion, friction, abrasion and fretting. The abrasion and fretting basically come from friction – unfortunately, they produce particles and debris on micro-, nano- and atomic scale. Therefore, the important interaction between particles and debris with tissue, biological fluids and the bone must be considered and taken into account [12].Figure 1:Artificial and natural hip (bearing) – typical components in a total hip replacement implant – a typical hip prosthesis with a combination of metallic and plastic components consists of the femoral stem, a femoral head and polymeric socket.Modular designs of the hip implant, where the stem and the ball (Figure 1) are made of two different materials, are common in practical applications. Modular orthopaedic hip implants are widely used in total hip arthroplasty due to their clinical flexibility. Another example of a common practical solution can be as follows: the ball is made of either highly polished Co-28Cr-6Mo alloy or ceramic sinters [e.g. alloyed alumina (Al2O3) with zirconia (ZrO2)], whereas stem is made of Co-based alloy [7], [8], [13], [14], [15], [16], [17]. Unfortunately, the modular type of hip prosthesis has revealed fretting corrosion and related problems, particularly at the taper interfaces. Also, a modular connection of the hip stem and head (Figure 1) made of different materials is susceptible to galvanic corrosion. Corrosion at the femoral stem/head taper interface may result from fretting due to micro motion, crevice at the taper mismatch and galvanic coupling of dissimilar materials or a combination of these three components.Titanium alloys have a tendency to stick. All titanium alloys have poor friction and abrasive properties. Therefore, sliding contact of these alloys should be avoided without modification of the surface layer by thermochemical diffusion treatment (nitriding and/or carbonitriding) [18], [19]. However, the effective case (and the total case) must be thick enough to be able to support the case layer. Also, the case layer should have a proper gradient-like distribution of microstructure, residual stress and micro hardness versus thickness to avoid the stress threshold or tension notch.Wear rates for UHMWPE mated against the Ti-6Al-4V alloy are significantly higher than that for Co-Cr-Mo alloys [6], [11]. Cobalt base alloys such as Co-Cr-Mo are the most commonly used metals for the current metal on polymer (UHMWPE) joint as friction couple in larger loaded bearings of hip implant components for heavier patients. Femoral components are also made from Co-Ni-Cr-Mo alloys. Important reports on cobalt alloys have recently been provided. It was about the easy diffusion of Co ions from hip implants to body fluids. In this way, concentration of cobalt in blood can reach a higher level than normal [20]. Although a majority of chemical elements of applied materials are building components of the human body, it is usually harmful if their concentrations in some tissues, bone or in body liquids are over a natural level. The higher wear rates of UHMWPE associated with its counterpart made of titanium alloy are related to the mechanical instability of the metal oxide layer under high loading. Mechanical stress and chemical activity cause micromechanical and chemical mechanisms of wear out called fretting corrosion. The fretting corrosion damage or mechanical fatigue is a major cause for failure of orthopaedic implants. Fretting can mechanically damage the passive layer and generally surface layers of all metals. The Ti-6Al-4V and other similar alloys are used for hip implants in the annealed state. These alloys in the annealed condition are relatively soft and have low strength, and the substrates of these alloys are not able to support the passive oxide layer exposed to heavy loading of working joints during use. Under a heavy load, the oxide layer collapses (cracks) and produces a flux of particles. Normal stresses usually are high enough to break down the surface passive layer causing oxide disruption. The passive oxide surface layer on femoral heads made of the titanium alloy results in an excessive wear of the couple metal-UHMWPE socket [7]. On the other hand, relatively soft alloy due to annealed microstructure with lower internal stored energy is required because of stable bioactivity and corrosion resistance due to lower internal stored energy. The bioactive surface of applied alloys for orthopaedic applications is very important for achieving improved biocompatibility. It is important that biomaterials will not introduce fibrous tissue barriers during contact with healthy bone. This permits the bone to grow closer to the surface of the implant and to fill in grooves or pores that have been deliberately introduced to a surface according to the Wolf law of bone growth to enable the device to become more firmly embedded.General requirements for orthopedic implantsArtificial solutions for human body disabilities must be based on the mechanical design and actually available materials, which meet special requirements. These are mechanical, biological and chemical requirements that must be fulfilled for them to be used for the human body in line with the principles of medicine describing the particular features of biomaterials. The above review of the literature reveals complex issues, which are related with medical, mechanical and material problems in hip implants. Biomaterials used for hip prostheses should exhibit the following properties [6], [7], [9], [10], [13], [21]:High biocompatibility (nontoxic and non-carcinogenic, chemically stable and corrosion resistant);Nonmagnetic;The surface femoral stem should be bioactive in terms of osteointegration;Adequate mechanical properties (Young modulus, yield strength, tensile strength, ductility, hardness, toughness, fatigue strength and abrasive wear);The femoral stem should be able to endure large and variable stresses in the human body environment relating to a large number of stress cycles, i.e. fatigue (e.g. during 10 years, assuming that the distance of 10 km in a day yields about 35×106 mechanical cycles, i.e. steps) with additional stress corrosion, stress-enhanced diffusion and even stress-induced phase transformation;Surface layer with gradient-like modulus of elasticity, i.e. with gradient-like porous distribution in surface layer;The surface of implants should be free of imperfection, such as tool marks, scratches, nicks, cracks, cavities, burrs and other defects;Surface texture as surface quality factor and surface geometry classification may vary from very smooth for bearings to rough for the femoral stem and socket. All this depends on the design, material, application and the area involved in the metallic implant (e.g. the fatigue crack initiation is significantly lowered when the surface is highly polished, whereas a rough surface is demanded for better osteointegration);Well-balanced chemistry and proper metallurgical conditions of the alloy, i.e. proper chemical, phase and microstructure composition, absence of segregation, nonmetallic inclusions and martensitic phase;Availability, with reasonable balance between quality, durability and price.Mechanical and materials requirements for orthopedic implantsOne of the goals in the development of new materials for hip implants was and still is an optimal ratio of the elasticity modulus of metallic, ceramic and polymer parts and the bone. The modulus is an important concern in the orthopaedic application of biomaterials in terms of the mechanical interaction between the implant and the bone. The stiffness of the material can be described by Young’s modulus of elasticity (E). The ratio of stress (s) to strain (e) in the elastic deformation region is known as the modulus of elasticity (E), or Young’s modulus. Young’s modulus is the tangent of slope angle of the σ=Ee strain line function in the elastic region of the tensile test. The constant of the proportion between the stress and strain in the linear elastic region during tension is known as Young’s modulus and defined as(1)Modulus of elasticity E=σx/εx,σx= Eεx(here E as stiffness),  εx= σx/E and εy= −νσx/E (here 1/E is susceptibility)$$\matrix {{{\text{Modulus of elasticity }}E = {\sigma _x}/{\varepsilon _x},} \hfill \\ \matrix {\sigma _x} = {\text{ }}E{\varepsilon _x}({\text{here }}E{\text{ as stiffness}}),{\text{ }}{\varepsilon _x} = {\text{ }}{\sigma _x}/E{\text{ and }} \hfill \cr {\varepsilon _y} = {\text{ }} - \nu {\sigma _x}/E{\text{ (here }}1/E{\text{ is susceptibility)}} \hfill \cr \hfill \\ } $$The modulus of elasticity (E) is essentially a measure of the material stiffness and is determined from the tensile test or by the elastic internal dumping method (internal friction) and by ultrasonic wave velocity measurements.The resistance to fatigue crack nucleation depends on stress, microstructure and modulus of elasticity according to the Paris law and Irvin and Weibull considerations [22], [23]. Fatigue cracks usually nucleate at the surface of the stressed metal, where the stresses are at a maximum taking into account surface residual stresses and the principle of superposition of any existing stress component. In contact loading, the Hertz and Bielajew point with the maximum reduced stress must be taken into account. Any design or manufacturing defects and tensile surfaces with residual stresses encourage nucleation and development of fatigue cracks. The fatigue resistance of the material can be expressed as the crack growth rate (da/dN) related to applied stress, strength (s, K) and Young modulus (E) of the material [23]:(2)dadN=CΔKm=Ca(KE)3.6$$\frac{{da}}{{dN}} = C\Delta {K^m} = {C_a}{\left( {\frac{K}{E}} \right)^{3.6}}$$whereda/dN is the crack growth rate [mm/cycle]C, Ca and m are the experimental constantsΔK=Δσ is the stress intensity factor during fatigue factorwhere sharp crack occurs, the existence of a critical stress for rapid crack propagation is given by the relation(3)σcrict=Eγ/πa$${\sigma _{{\text{crict}}}} = \sqrt {E\gamma /\pi a} $$where σcrit is the critical tensile stressE is Young’s modulusγ is the energy per unit area of new surfacesa is the crack length for a surface crack and 2a=for an internal crackThe stress from the applied loading must be lower than the tensile strength of the material (Table 1). If the tensile strength of the material increases, the resistance to fatigue also increases by the crack growth rate decrease [Eq. (2)]. If Young’s modulus is larger, the crack growth rate is smaller and the critical stress to develop crack must be larger [Eq. (3)].Table 1:Young modulus and mechanical properties of the bone and selected biomaterials.Properties/materialYoung’s modulus E, GPaPoisson’s ratioTensile ultimate strength Rm, MPaTensile yield strength Re, MPaRe/RmBending strength Rg, MPaUltrahigh molecular weight polyethylene (UHMWPE)55–1700.42>30>20––Bone cement3.80.460550.9–Compact bone30–120––160Ti CP Grade 4/annealed1050.37560–880380–5200.7340Ti-6Al-4V/annealed1150.41>890>7300.9650Cobalt alloys (Co-28Cr-6Mo)/annealed2300.36>600>5000.6–Bio-steels (see Chap. austenitic steels)2100.29480–900180–6000.6250Reciprocal relation appears when there are differences in an elastic module of the bone and the implant. Significant differences in the elastic moduli of the bone and the implant may lead to stress concentrations at their interface, i.e. the stress threshold. An insufficient load transfer from the artificial implant to the adjacent natural bone may result in bone desorption and loss of the bone-implant joint, and finally, the whole hip implant can be rejected. The low modulus of alloys reduces the mismatch between the modulus of the implant and that of the bone, which in turn reduces the so-called “stress shielding phenomenon”, which can be responsible for potentially damaging resorption of bone on the inner surface of the natural femur. Stress shielding is related to the difference in flexibility/susceptibility or stiffness between natural bone and the implant alloy. The lower modulus of titanium alloys is a positive factor in reducing bone resorption. However, the modulus of elasticity should not be too low. An implant made from alloy with the too low Young’s modulus of elasticity will have low resistance to nucleation of fatigue cracks and larger cracking velocity [Eq. (3)]. From this point of view the modulus of elasticity (E) should be in the optimum range owing to the selection of proper materials and/or proper production technology.Recent attempts to minimize the modulus of elasticity and improve biocompatibility have led to the introduction of some metastable β titanium alloys with elastic modulus values as low as about 70–80 GPa [24], [25], [26]. However, there has been, and still is, concern about too large elastic modulus of implant alloys compared to that of compact bone (30 GPa). The high difference in the modulus of elasticity of the alloy and of the bone causes stress incompatibility by stress, stress concentration and stress threshold. The modulus of elasticity should be optimum which assures better transfer of stresses between femoral stem with flange and a porous coating of alloy with gradient porosity, which means a gradient-like distribution of Young’s modulus. The modulus of elasticity depends on the degree of porosity. The modulus of elasticity decreases with increasing the porosity of coating from the implant surface to the bone.Amongst metallic biomaterials, the elastic modulus of Ti-6Al-4V alloy (E=110 GPa) is much lower than that of 316L (X3CrNiMnMo18-14-5-3) stainless steel (E=210 GPa) or Co-Cr-Mo alloy (E=230 GPa). Other mechanical properties like hardness, ultimate tensile strength (UTS), Rm, yield strength (R0.2), bending strength (Rb), torsion strength, toughness and fatigue strength of the alloys considered should be also measured and analyzed (Table 1).Mechanical properties depend on the microstructure which is characterized by grain size, grain crystallographic orientation, crystal lattice defect density and phase composition. The last one is formed by phase transformation due to heat, thermochemical and thermomechanical treatments. Metal forming, i.e. cold and/or heat plastic deformation, are important technological operations which influence lattice defects, grain size and their orientation. Segregation of alloying elements can result in the formation of non-uniform phases and microstructure composition with different crystallographic structures, lowered corrosion resistance and fatigue strength and increased residual stresses and pitting corrosion. Agglomeration or large amounts of other phases can reduce the fatigue life of implants. Segregation of non-uniform clusters acts as stress threshold in the microstructure. An alloy with segregated alloying elements is sensitive to nucleation of fatigue cracks. The fatigue or stress corrosion damage and abrasive wear are the major causes for failure of orthopaedic implants. The fatigue strength yield point ratio of applied biomaterials is small and ranges between 0.3 and 0.7 with additional large safety coefficient which makes the calculations of mechanical strength and total reduced stress rather complex and difficult.Biomedical, mechanical and material engineering of hip implantsThe natural constitution of the hip joint, loading stresses, residual stresses and corrosion, mechanical design of human bearing and the materials which are available define problems to solve for artificial hip implants. Mechanical, chemical and physical properties as well as biocompatibility requirements define suitable solutions for this complex problem. Typical components of hip implants are shown in Figure 1.As shown in Figure 1, a typical hip prosthesis consists of a femoral stem, a femoral head and socket. The femoral stem is usually manufactured from Co-Cr-Mo or Co-Ni-Cr-Mo alloys or from titanium base alloys (Ti-6Al-4V or new one Ti-Nb-Zr). The head is made of highly polished Co-Cr-Mo alloy, of Al2O3 sintered ceramic composite or of stabilized ZrO2 with Al2O3 and yttrium oxide additives [10]. The metal-backed shell and the screws are made of titanium alloy Ti-6Al-4V or CP titanium (Tables 1 and 2). The socket (cup) is made of UHMWPE usually with metallic shell. Modular designs, where the stem and head are made of two materials, are common [11]. However, when two different metallic alloys are applied, problems with galvanic corrosion should be considered and taken into account.Table 2:Chemical composition of commercially pure titanium grades in wt.% according to ASTM F 67.DesignationTiC (max)Fe (max)N (max)O (max)H (max)OtherASTM Grade 199.50.080.200.030.180.15–ASTM Grade 299.20.080.500.030.200.15–ASTM Grade 399.10.080.250.050.300.15–ASTM Grade 499.00.080.500.050.400.15–ASTM Grade 799.20.100.300.030.250.150.2 PdProduction technologyMechanical and biomedical properties depend on chemical, phase and microstructure composition. The technological processes which are applied influence characteristics of the above materials. Therefore, knowledge about the relationship between particular technological operations and the properties considered is an important key to understanding the technology of production of implant parts and their expected features. Generally, metallic parts are produced by casting, metal forming (forging, rolling and pressing), powder metallurgy (compacting and sintering) and various types of heat treatments and machining. The most recent manufacturing method, i.e. “3D printing” as an additive technology, is under development. Commonly, metallic parts of hip implants are manufactured by forging. Investment casting of hip stem is used as manufacturing method which is an alternative to forging. The main disadvantages of parts produced by investment cast are lower mechanical properties than those of forged parts. Forged hip stems have a higher yield strength, tensile strength and fatigue strength than those made by a cast. This is due to the proper microstructure of alloy resulting from large plastic deformation during forging. The microstructure of cast alloys, called primary microstructure, is generally worse due to non-uniformity like segregation of alloying elements and different grains size in different places. Mechanical properties of forged components are better than those produced by casting or 3D printing technology. The femoral stem is usually forged as α+β titanium phases at a temperature of 955°C. Forged stem parts are then annealed at recrystallization temperature. The microstructure is either mill annealed or fully equiaxed depending on the parameters of annealing according to the principle: the higher the temperature of annealing, the larger the grain size and the smaller the mechanical properties. Mechanical properties of this alloy can be controlled over significant range through hot and/or cold working and heat treatment.The compacting and sintering technology is used to produce ceramic components. The following technological operations are heat or thermochemical treatments including stress relaxation, annealing, super saturation, ageing, carburizing and nitriding. Grinding and polishing as the final machining are applied to obtain the final shape and quality of a surface. The introduction of compressive stresses to the surface layer by thermochemical treatment (for example, by nitriding or carbonitriding) or shot penning increases resistance to fatigue cracking. The majority of all the above considerations relate to metallic components of hip implants; however, the contemporary modular structure of hip implants still needs to be developed in the field of design, manufacturing technology and materials. The advanced metallic, ceramic, composite and polymer materials are an important key to the progress of hip prosthesis and of biomedical engineering in general.Introduction of titanium and titanium alloys for biomedical engineeringThe selection of titanium and titanium alloys for hip implants is determined by a combination of most favourable characteristics including high strength to weight ratio, outstanding resistance to corrosion by body fluids, high biocompatibility, low density and the capacity for joining with bone, i.e. osteointegration. Titanium alloys have outstanding resistance to corrosion by body fluids, which is superior to that of stainless steels. Although titanium is highly reactive to oxygen, the creation of passive oxygen layers and solid interstitial solution exhibits excellent corrosion resistance in the human body environment. Sliding contact should be avoided without modifying the surface by thermochemical diffusion treatment (such as nitriding see, e.g. Figure 2).Figure 2:X-ray diffraction pattern of nitrided Ti-6Al-4V alloy with Bragg-Brentana and grazing angle (α) diffraction geometry for comparison of phase composition versus X-ray depth of penetration (z), (for BB z=5–18 μm; for α=2, z=1.1 μm; for α=4, z=2.4 μm; and for α=9, z=4.8 μm) [27].Titanium exists in two allotropic crystallographic forms: Tiα, which has the hexagonal closed packed crystallographic structure, and Tiβ form, which has the body cubic (BCC) structure. In pure titanium, the α phase is stable up to 882°C. Above 882°C the α→β transit temperature appears. The hexagonal α phase of pure Ti is transformed in heating to the BCC β phase at a constant temperature of 882°C. This phase transformation temperature depends on the chemical composition of titanium alloys [28]. Commercially pure titanium which is unalloyed ranges in purity from 99.5 to 99.0% Ti (Table 1). The main elements in unalloyed titanium are iron and interstitial elements like carbon, oxygen, nitrogen and hydrogen [2], [4], [19], [21]. The chemical compositions of the principal grades of titanium are listed in Table 1. Commercially pure titanium can be considered as an α phase alloy in which the oxygen content determines the grade and strength. Titanium sponge due to metallurgical technology contains oxygen at a certain level, but its amount is adjusted to modify the strength of commercially pure titanium. In this case, oxygen can be treated as an “alloying element”. Carbon, nitrogen and hydrogen are present as metallurgical impurities in titanium. Oxygen within solution limits can be used as an interstitial strengthening element. Hydrogen is readily absorbed by titanium alloys and is detrimental because it causes so-called hydrogen embrittlement by brittle needle-like precipitation of titanium hydrides (TiH2). Oxygen, nitrogen and hydrogen, on the other hand, can improve biocompatibility. Nitrogen, carbon, oxygen, boron and hydrogen form interstitial solid solutions because of the large difference between the atom size of titanium and these elements. There is a reasonable difference in solubility of these elements in α and β titanium [28]. Titanium and its alloys react with several interstitial elements including gaseous oxygen and nitrogen, and such reactions occur at a temperature well below the respective melting points. This chemical affinity is used in the thermochemical treatment of titanium alloys.Commercially pure titanium of ASTM grades have low strength but higher corrosion resistance than titanium alloys. The addition of 0.2% Pd to commercially pure titanium improves its strength and corrosion resistance in active reducing media and is designated as Grade 7 (Table 2).Although interstitial elements increase the strength of titanium, they are detrimental to toughness as measured by the notch impact test. Therefore, when high toughness is desired for certain applications, the alloy should be produced with extra low interstitials (ELI). These alloys are referred to as ELI alloys [27]. Strain hardening of titanium alloys is strongly sensitive to the rate of deformation. Increasing rate of deformation increases significantly the strain hardening rate. In order to obtain a large deformation of titanium alloys, the process is carried out in hydraulic presses with a slow deformation rate. A slower deformation rate produces larger deformation of titanium alloys without cracking. Plastic deformation is realized with a dislocation slide mechanism suitable to the crystallographic lattice of both α and β phases [29]. The twining system of deformation appears in both titanium phases as well.The titanium α/β alloys are most often used in the annealed condition. Their microstructure and mechanical properties may differ depending on whether or not prior plastic deformation was carried out above or below the α/β phase transformation temperature. A typical forging temperature of α/β Ti-6Al-4V is 955°C. An inert gas, e.g. argon can be introduced into one part of the mould cavity. During the heat treatment, the high vacuum atmosphere is used, 10–5 torr). Above 535°C, titanium absorbs oxygen and forms oxide compounds (TiO, Ti2O3 and TiO2) and a solid solution as a subsurface layer. All of them can improve mechanical properties from one side and nucleate surface and brittle cracks from another side. Titanium also absorbs hydrogen, and it is necessary to ensure that furnace atmospheres are hydrogen free during melting, heat and thermochemical surface treatments. Titanium alloys are particularly susceptible to galling, i.e. wear due to friction, during hot or cold working, which causes surface damage. Therefore, proper lubricants should be applied during thermomechanical cold or hot working. During these processes, graphite or molybdenum disulphide is used, whereas glass powder may be used for a more severe process like hot extrusion [27].Physical metallurgy and surface layers of biocompatible titanium alloys for hip implantsTitanium alloys are sensitive to contact in friction couple due to sticking or galling, which may lead to fretting. For this reason, titanium alloys have low resistance to fretting. Fretting occurs when two surfaces are in contact and small amplitude relative oscillation motion is present. Fretting damage in titanium alloys can significantly reduce fatigue life. Fretting is a major concern for hip implants in joints between femoral stems and heads.Fretting corrosion is a serious problem in implants because it is the main mechanism of gradual release of metal ion to different tissues in the body. Fretting corrosion and abrasive wear cause pain and excessive accumulation of wear particles which results from prosthesis losses. The extensive release of metal ions from hip prosthesis can result in adverse biological reactions [12], [19].Two primarily problems arise when attempting to coat titanium alloy with TiN, TiC and TiCN layers. The thermal expansion coefficients of the coating and of the titanium substrate are different, and as a result, large thermal stresses are generated at the interface during the coating deposition process [30]. Such coating also has a different microstructure, crystallographic structure, different mechanical properties (strength, hardness, modulus of elasticity etc). The coating layer of the femoral head should not be too hard. Under heavy load and residual stress, the coated layer will be cracking. These stresses can lead to cracks at the interface coating/substrate. In addition, chemical reactions between the coating and titanium alloy can weaken the metal in the vicinity of the interface, reducing strength and corrosion resistance of the coating system. Most advanced coatings were elaborated with the application of hydroxyapatite and some other ceramic compounds [19], [31].Therefore, diffusion thermochemical treatment is a better solution in order to produce hard, highly resistant to wear and gradient like mechanical properties versus depth from the surface. Here is an example of a thermochemical ion nitriding of Ti-6Al-4V alloy (Figure 2). Very clean nitrogen (N2) was introduced to vacuum furnace chamber. Under high voltage, electrical discharge occurs and the nitrogen is dissociated, ionized in the form of glow discharge which produces single ions (in statu nascendi) according to the following reaction:N2→N2++N2++4e${{\text{N}}_2} \to {{\text{N}}^{2 + }} + {{\text{N}}^{2 + }} + 4{\text{e}}$Only an active nitrogen in “statu nascendi” can diffuse to the metal surface layer of metal. The diffusion ion nitriding treatment (plasma nitriding) was carried out in controlled atmosphere (containing a mixture of nitrogen ions and argon) at a temperature of 700°C. The phase composition varies versus thickness of the surface layer, and some level of residual stresses was measured (about −522 MPa) with GID-sin2ψ method [30]. The qualitative X-ray phase analysis showed the dominating phase component as Tiα, TiN and Ti2N (Figure 2). The two latest nitrides appear in a thin surface layer of a nitrided Ti-6Al-4V alloy with the plasma thermochemical process. The investigated surface layer was exposed with X-ray grazing angle (α) diffraction geometry with incident beam angle α=2 and α=4°. The applied X-ray diffraction geometries are destined for thin surface layer analysis with different thicknesses. This means that very thin surface layer thickness of 2 and 4 μm consists of TiN and Ti2N, and deeper layers as a substrate under them consist of hexagonal phase Tiα(N) as a solid solution with a gradient-like nitrogen distribution (Figure 2). Similar experiments and examinations carried out on CP4 titanium [18] under a little bit different conditions confirmed phase composition and compressive residual stresses.Classification of titanium alloysTitanium alloys are classified into categories according to the microstructure and phase composition (structure) retained at room temperature after solidification and other metallurgical processes. Both characterizations depend on heat and various kinds of thermomechanical treatments. According to the phase composition (as structure composition), titanium alloys can be divided intoα alloysnear α alloysα+β alloysβ alloysnear β alloysmetastable β alloysThe most known Ti-6Al-4V titanium alloy contains 6 wt.% aluminium and 4 wt.% vanadium. It was invented in 1946 in the USA as a most advanced alloy for aeroplane and cosmos industry applications. Soon afterwards, this alloy was experimentally introduced as a good biocompatible alloy for medical application. Gradually, aluminium and later vanadium were substituted by elements more friendly for the human body (Table 3). A new generation of titanium alloys with niobium was introduced in 1979 [5], [25], [26], [32], [33], [34]. New β or near β titanium alloys contain biocompatible alloying elements: Nb, Zr and Ta. They have a lower modulus of elasticity as compared to α+β alloys. The β alloys have excellent formability, cold rolling capabilities, good ductility and formability. The main disadvantage of β alloys in comparison with α+β alloys is their higher density. Most recently developed alloys based on the Ti-Nb-Zr-Ta alloying elements appeared as β alloy has the lowest modulus of elasticity, E=70 GPa [3], [24].Table 3:Selected experimental and applied titanium alloys for implant applications.Alloy designation and typeDeveloped inα+β alloysTi-6Al-4VUSATi-6Al-4VELIUSATi-6Al-7NbSwitzerlandTi-5Al-2.5FeGermanyTi-3Al-2.5VUSATi-15Zr-4Nb-2Ta-0.2PdJapanTi-5Al-3Mo-4ZrJapanTi-15Sn-4Nb-2TaJapanβ alloysTi-15Mo-5Zr-3AlJapanTi-29Nb-13Ta-4.5ZrJapanTi-13Nb-13ZraUSATi-12Mo-6Zr-2FeUSATi-16Nb-10HfUSATi-35Nb-7Zr-5TaUSATi-15Mo-2.8Nb-0.2Si-0.26OUSAaThe only alloy with homogeneous microstructure is single-phase β. Alloy with segregation of Nb and Zr is not acceptable. Segregation of alloying elements (Nb and Zr) causes not homogeneous microstructure sensitive to corrosion and fracture.Alloy Ti-13Nb-13Zr was originally developed for medical implant applications [24]. This alloy combines a high biocompatibility, high strength, lower elastic modulus than alloy Ti-6Al-4V, excellent hot and cold workability and superior corrosion resistance. The Ti-13Nb-13Zr alloy is a metastable beta titanium alloy developed for use in biomedical implants that combines a high biocompatibility, superior corrosion resistance, high strength, excellent hot and cold workability and lower elastic modulus. Mechanical properties of this alloy can be controlled over significant range through hot working, cold working and heat treatment. The mechanical properties of Ti-13Nb-13Zr can vary by thermomechanical treatment. This alloy contains a high percentage of high-density alloying elements. Because of the high density of zirconium and niobium, there is a problem of segregation during melting and solidification of alloys containing these elements.Also, because of the high density of niobium and zirconium, special techniques of mixing liquid metal during melting and casting process are required to prevent segregation of alloying elements. Niobium and zirconium have larger atomic diameter than titanium. Special thermochemical treatment which provides point defects and dislocations should be applied to increase the diffusion rate of alloying elements in order to obtain the proper response during homogenization treatment. Segregation of alloying elements cannot be tolerated in metallic implants. Segregation of alloying elements decreases resistance to corrosion and fatigue crack initiation, lowers resistance to an impact load and fracture resistance and decreases biocompatibility. The newest development of titanium alloys is directed to Ti-Au composition with an expectation of better mechanical and biomedical properties. Another titanium base alloy with nickel, nitinol, is also an important biomaterial [35], [36].Cobalt base alloys for hip implantsCobalt-chromium alloys have good resistance to pitting and crevice corrosion in the human body. The main cobalt alloys (Table 4) used for hip implants areASTM F75, (Co-28Cr-6Mo) casting alloyASTM F90, (Co-20Cr-15W-10Ni) wrought alloyASTM F999, (Co-28Cr-6Mo) thermomechanically processedASTM F562, Co-35Ni-20Cr-10Mo wrought alloyASTM F 1058 (Co-20Cr-15, 5 Ni-7Mo-Fe) wrought alloyMP35N, (35Co-35Ni-20Cr-10Mo) thermomechanically processedTable 4:Chemical composition in wt.% of selected cobalt base alloys used for hip implants (Co balance).ASTM designationUNS no.CrMoNiFeCSiMnWPSOtherF75R3007527.0–30.05.0–7.01.00.750.351.01.00.200.0200.010.25 N, 0.30 AlF90R306519.0–21.0–9.0–113.0 max0.05–0.150.400.4014.0–16.00.0400.03–F562R3056319.0–21.09.0–10.533.0–37.01.0 max0.025 max0.150.15–0.0150.011.0 TiF563R3056318.0–22.03.0–4.015.0–25.04.0–6.00.050.501.03.0–4.0–0.010.50–3.50 TiF799R3153726.0–30.05.0–7.01.00.750.351.01.0–––0.25 NBiophysics and physical metallurgy of cobalt alloysCobalt is beneficial for humans because it is a component of vitamin B12 (cobalamin), which is essential for human health. However, too high concentrations of cobalt may be toxic to health. A cobalt-based alloy was first used as a material for hip implants in the USA in 1936. At temperatures below 417°C cobalt exhibits a hexagonal close-packed structure. Between 417°C and its melting point of 1493°C, cobalt has a face-centred cubic (FCC) structure. The melting point of cobalt is 1768°C. The elastic modulus of cobalt is about 210 GPa in tension and about 183 GPa in compression. Cobalt has a density of 8.90 g/cm3. Typically, the microstructure of cobalt-based alloys consists of FCC as (γ) phase as matrix and various types of carbides. A strengthening in cobalt-based alloys is obtained primarily through a combination of solid solution strengthening (by Cr, Mo and Fe atoms) and carbide precipitation. Strength and toughness depend on the amount of solved foreign atoms and amount, size and distributions of carbides. A fine dispersion of carbides contributes significantly to the strength of these alloys. Dislocation movement is strongly impeded by carbide precipitations. Carbide precipitations are barriers for dislocation movements. This carbon content in these alloys has a significant effect on the number of carbides. In general, three main types of carbides can be present in the microstructure of cobalt-based alloys [8], [14], [15], [16], [17]:M23C6 carbides, where “M” is mostly chromium but can be substituted by tungsten and molybdenum;MC carbides, where M stands for the reactive metals tantalum, titanium, zirconium and niobium;M6C carbides, where “M” stands for tungsten or molybdenum; these carbides form when the tungsten or molybdenum content exceeds about 5 at.%.Components made of cobalt alloys are designed as perfect spherical head of hip bearing so that the metal always articulates against a low friction plastic (e.g. UHMWPE or high cross-linking polyethylene), which provides smooth relative movement, small friction and minimum abrasive wear. The recent reports on cobalt alloys applied for hip implants show an easy diffusion of Co ions from the hip implant to body fluids. This type of implementation with cobalt alloy implants caused excessive concentration of Co in blood [20]. Therefore, some limitations in this type of application have been prepared.Austenitic stainless steelsThe main components of stainless steels are iron, nickel and chromium. Generally, stainless steels are divided into three groups: ferritic, martensitic (tool steels) and austenitic. Only the last type of stainless steels called bio-steels is applied for a hip prosthesis for short-term use. Austenitic stainless steels were first used as a material for hip implants in the USA in 1926.Nevertheless, iron ions are important components of blood; other chemical components are rather harmful to human body.Austenitic stainless steels are used for implant applications because they are relatively inexpensive. Passivation of stainless steel implants is enhanced by nitric acid. Austenitic stainless steels are not sufficiently corrosion resistant for long-term use as an implant material. Pitting corrosion in austenitic stainless steels is usually caused by chlorine ions and starts at nonmetallic inclusions (sulphur inclusions) and other heterogenic places. Molybdenum is added to stainless steels usually in amounts of 2–3% to increase resistance to pitting corrosion in saline environments. Nitrogen-strengthened alloys are used for the bone plate and bone screws.Austenitic stainless steels have excellent formability, and their response to deformation depends on the nickel content. Formability increases with an increase in the content of nickel. Magnetic alloys should not be used in the human body because they could become dislodged in the strong magnetic field of magnetic resonance imaging. Pitting corrosion in the stainless steels occurs when the protective film is broken, exposing the steel to Cl− ions. Austenitic stainless steels for hip implants can be divided intolow carbon stainless steelsnitrogen-strengthened stainless steelsnickel-free stainless steelsExamples of major ASTM austenitic stainless steels for implant materials:Low carbon stainless steels: ASTM F138/139 316 L, (21Cr-10Ni-3Mn-2.5 Mo). This steel contains 0.25–0.50% N.Nitrogen-strengthened stainless steels: ASTM 1314 (22Cr-12, 5Ni-5Mn-2.5Mo-0.30 N). This steel contains 0.20–0.40% N.Nickel free stainless steel and nitrogen-strengthened steels: ASTM 1314 F2229 (23Mn-21 Cr-1Mo –0.95N).The substitution of nickel atoms is important due to its harmful activity in the human body. Because of that, Ni-free stainless steels were elaborated (Table 5).Table 5:Chemical compositions (in wt.%) of austenitic stainless steels used for surgical implants according to the ASTM specifications (Fe balance).aASTM designationCCrNiMnMoCuNPSOtherF1380.0317.00–19.0013.0–15.002.002.25–3.000.500.100.0250.010–F13140.0320.50–23.5011.50–13.504.00–6.002.00–3.000.500.20–0.400.0250.0100.10–0.30 Nb; 0.10–0.30 VF15860.0819.50–22.009.00–11.002.00–4.252.00–3.000.250.25–0.500.250.0100.25–0.80 NbF22290.0819.00–23.000.1021.00–24.000.50–1.500.250.90 min0.030.0100.95 NaSingle values are maximum values unless otherwise indicated.SummaryThe review, as presented above, i.e. titanium and cobalt alloys, bio-steels, their specific treatments, surface layers and coatings on alloys used for hip prosthetic can be summarized as follows:The synergy of some branches of science and technology, i.e. medicine (surgery), mechanical design and material science and technology has led to progress and achievements in hip implants.All of the chemical elements of the materials applied in implants are building components of the human body, but some of them are even poisonous and others, like aluminium, vanadium and nickel, are classified as toxic elements.Biomedical implant reliability depends largely on the corrosion, abrasive wear and fatigue strength/resistance of the materials used in production technology.The capacity for joining implant with the bone, i.e. osteointegration, which is direct bone anchorage to an implant into bone with no fibrous capsule interface, makes it possible to transmit the loading forces directly to the bone.All the beneficial passive layers on metals can be mechanically damaged by the fretting of metal against metal.Titanium alloys are used for hip implants due to excellent combination of mechanical properties, low density, acceptable tissue tolerance, high strength to weight ratio, outstanding resistance to corrosion by body fluids, high biocompatibility, low density and nonmagnetic properties.Titanium induces the formation of a fibrous tissue barrier when placed in contact with a healthy bone. This permits the bone to grow closely on the surface of the femoral stem and fill pores that have been deliberately introduced to facilitate bone ingrowths.Contaminations of titanium alloys by interstitial elements such as hydrogen and oxygen during fabrication (melting, thermomechanical treatment and thermochemical surface hardening) must be avoided because these elements have an embrittling effect on titanium.In the presence of hydrogen, titanium hydrides (TiH2) may form into long thin needles in titanium alloys and these hydrides decrease resistance to fracture.Austenitic stainless steel, particularly nickel-free stainless steel with nitrogen strengthened like ASTM 1314 F2229 (23Mn-21Cr-1Mo-0.95N), can be used for some elements of hip implants.Cobalt-chromium alloys are the most commonly used metals for current metal on polymer implants. A typical microstructure of cobalt base alloys consists of FCC (γ) phase matrix and carbides of various types.Hydroxyapatite is widely used as bioactive porous coating of the femoral stem and the metallic shell in hip implants.A special type of polyethylene, i.e. special grade UHMWPE with additional cross linking bonding, is widely used (as socket) in hip implants as the load-bearing material (in hip implants with combination metal on polymer).The final decision about the type and material of hip implants made in the hospital by medics is supported by many additional factors like accessibility, availability of the material, the patient’s weight and age, etc.ConclusionsThe synergy of medicine, biochemistry, biophysics, mechanical engineering, materials science and informatics succeeds in achieving real progress in orthopaedic surgery of hip joint implants.Although the majority of chemical elements of the materials applied to make implants are building components of the human body, it is usually harmful if their concentrations in some tissues, bone or in body liquids are over the natural level.The fatigue and stress corrosion damage are the major causes for the failure of orthopaedic implants. The fatigue strength and yield point ratio of biomaterials applied are small and range between 0.3 and 0.7, which, with additional large safety coefficient, cause mechanical strength and total reduced stress calculations to be rather complex and difficult.The majority of parts of hip joint implants are made of metallic materials due to heavy and cyclic loaded bearings. Additionally, they work in the bioactive environment. Titanium alloys are most advanced materials in this type of application; however, Co-Cr-Mo alloys and austenitic stainless like bio-steels mated with appropriate metallic alloys are the main materials for components of hip implants.The natural constitution of the hip joint; the stresses resulting from loading, residual stresses and corrosion; mechanical design of human bearing; and accessible materials define problems to solve for artificial hip implants.The contemporary modular structure of hip implants still needs to be developed in the field of design, manufacturing technology and materials. The advanced metallic, ceramic, composite and polymer materials are an important key to the progress of hip prosthesis and of biomedical engineering.NoteThe information in this article is for informational and educational purposes and is not meant as medical advice or recommendation. Only a qualified orthopaedic surgeon can determine which material implant system is best for an individual person. There are many factors that the surgeon uses when recommending hip implants (state of health, weight, age, life activity, anatomy etc.). Any questions and concerns regarding specific type of materials and implants should be discussed with the professional orthopaedic surgeon. Every patient’s case is unique, and each patient should follow specific instructions of his or her doctor. The information in this article does not replace orthopaedic doctor’s specific instructions.Acknowledgments:The authors would like to thank MSc Kamil Nawojowski for his support.Author contributions: All the authors have accepted responsibility for the entire content of this submitted manuscript and approved submission.Research funding: The results of research and this paper were supported by the Faculty of Metals Engineering and Industrial Computer Science, Department of Physical and Powder Metallurgy of AGH-University of Science and Technology in Krakow under project no 11.11.110.299.Employment or leadership: None declared.Honorarium: None declared.Competing interests: The funding organization(s) played no role in the study design; in the collection, analysis, and interpretation of data; in the writing of the report; or in the decision to submit the report for publication.References1.Niinomi M. Recent titanium R&D for biomedical applications in Japan. 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Biomaterials for hip implants – important considerations relating to the choice of materials

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Abstract

IntroductionThe first hip joint replacement took place in 1926 in the USA. Hip joint replacement is a major achievement in orthopaedic surgery in the 20th century. Scientific research and experiments with a large number of human body part implants are carried out. Medicine, mechanical engineering, chemistry, biology, physics and material science have created particular divisions and branches intended to investigate implants.The synergy of medicine, biochemistry, biophysics, mechanical engineering, materials science and informatics succeeds in development of real progress in orthopaedic surgery of hip joint implants. Recently, several millions of hip joint replacements are made per year in the USA. Similar progress takes place in many other countries. Among others, the quality of the materials used and the material science, in general, are involved in achievements in this interdisciplinary field. Therefore, this scientific field with respect to hip joint implants is discussed and considered in this article.Most parts of hip joint implants are made of metallic materials due to heavy and cyclic load bearing. Additionally, they work in a bioactive environment. Titanium alloys are the most advanced materials in this type of application; however, Co-Cr-Mo alloys and austenitic stainless steels as bio-steels mated with appropriate metallic alloys, ceramics and polymers are the main materials of hip implant components. Femoral heads and sockets (cups) can also be made from advanced ceramic materials. These statements are based on experience and scientific biomedical investigations [1], [2], [3], [4], [5], [6], [7], [8], [9], [10]. Additionally, ultrahigh molecular weight polyethylene (UHMWPE), high-density cross linking polyethylene and polieteroeteroketon are applied to produce the circular bearing.Metallic alloys are crystalline materials since their properties strongly depend on the type of the crystal lattice. The specific space atomic arrangement defines the crystal lattice, which in turn relates to the particular properties and their anisotropy. The chemical and phase composition of an alloy as the main characteristic has a major influence on the crystal lattice, grain size, lattice defects and crystalline texture. The microstructure revealed with microscopic investigation provides information about the shape and size of grains and precipitations, which is the third important characteristic of biomaterials. All these properties of biomaterials are formed during technological operations of production like the metallurgical process, solidification, sintering, metal forming (rolling and forging), machining, heat treatment and coating deposition.All these technological processes are optimized in order to achieve the optimal structure and microstructure coupled with the expected mechanical, chemical and biological properties. The metallic components usually undergo additional surface treatments including coating and alloyed surface layers deposition.Recently, titanium alloys are most often used in biomaterial applications due to an excellent combination of properties. The titanium alloys are characterized by good mechanical properties, low density, an acceptable tissue tolerance, high strength to weight ratio, outstanding resistance to corrosion by body fluids, high biocompatibility, being nonmagnetic and the capacity for joining with bone, i.e. osteointegration [1], [4], [11]. Amongst others, it is important that titanium is a material that will not introduce fibrous tissue barriers in contact with a healthy bone. Another advantage of titanium is the fact that the fatigue properties of the load bearing device are not reduced through contact with body fluids containing aggressive chlorine ions. The biological, chemical and mechanical contacts of foreign material with the bone, body fluids and tissues are most important characteristics of the implant, although the interaction between mechanical parts of a complex and modular implant is an additional serious area of scientific investigations. The components of modular implants (Figure 1) are exposed to degradation processes as fatigue loading, contact fatigue, corrosion, friction, abrasion and fretting. The abrasion and fretting basically come from friction – unfortunately, they produce particles and debris on micro-, nano- and atomic scale. Therefore, the important interaction between particles and debris with tissue, biological fluids and the bone must be considered and taken into account [12].Figure 1:Artificial and natural hip (bearing) – typical components in a total hip replacement implant – a typical hip prosthesis with a combination of metallic and plastic components consists of the femoral stem, a femoral head and polymeric socket.Modular designs of the hip implant, where the stem and the ball (Figure 1) are made of two different materials, are common in practical applications. Modular orthopaedic hip implants are widely used in total hip arthroplasty due to their clinical flexibility. Another example of a common practical solution can be as follows: the ball is made of either highly polished Co-28Cr-6Mo alloy or ceramic sinters [e.g. alloyed alumina (Al2O3) with zirconia (ZrO2)], whereas stem is made of Co-based alloy [7], [8], [13], [14], [15], [16], [17]. Unfortunately, the modular type of hip prosthesis has revealed fretting corrosion and related problems, particularly at the taper interfaces. Also, a modular connection of the hip stem and head (Figure 1) made of different materials is susceptible to galvanic corrosion. Corrosion at the femoral stem/head taper interface may result from fretting due to micro motion, crevice at the taper mismatch and galvanic coupling of dissimilar materials or a combination of these three components.Titanium alloys have a tendency to stick. All titanium alloys have poor friction and abrasive properties. Therefore, sliding contact of these alloys should be avoided without modification of the surface layer by thermochemical diffusion treatment (nitriding and/or carbonitriding) [18], [19]. However, the effective case (and the total case) must be thick enough to be able to support the case layer. Also, the case layer should have a proper gradient-like distribution of microstructure, residual stress and micro hardness versus thickness to avoid the stress threshold or tension notch.Wear rates for UHMWPE mated against the Ti-6Al-4V alloy are significantly higher than that for Co-Cr-Mo alloys [6], [11]. Cobalt base alloys such as Co-Cr-Mo are the most commonly used metals for the current metal on polymer (UHMWPE) joint as friction couple in larger loaded bearings of hip implant components for heavier patients. Femoral components are also made from Co-Ni-Cr-Mo alloys. Important reports on cobalt alloys have recently been provided. It was about the easy diffusion of Co ions from hip implants to body fluids. In this way, concentration of cobalt in blood can reach a higher level than normal [20]. Although a majority of chemical elements of applied materials are building components of the human body, it is usually harmful if their concentrations in some tissues, bone or in body liquids are over a natural level. The higher wear rates of UHMWPE associated with its counterpart made of titanium alloy are related to the mechanical instability of the metal oxide layer under high loading. Mechanical stress and chemical activity cause micromechanical and chemical mechanisms of wear out called fretting corrosion. The fretting corrosion damage or mechanical fatigue is a major cause for failure of orthopaedic implants. Fretting can mechanically damage the passive layer and generally surface layers of all metals. The Ti-6Al-4V and other similar alloys are used for hip implants in the annealed state. These alloys in the annealed condition are relatively soft and have low strength, and the substrates of these alloys are not able to support the passive oxide layer exposed to heavy loading of working joints during use. Under a heavy load, the oxide layer collapses (cracks) and produces a flux of particles. Normal stresses usually are high enough to break down the surface passive layer causing oxide disruption. The passive oxide surface layer on femoral heads made of the titanium alloy results in an excessive wear of the couple metal-UHMWPE socket [7]. On the other hand, relatively soft alloy due to annealed microstructure with lower internal stored energy is required because of stable bioactivity and corrosion resistance due to lower internal stored energy. The bioactive surface of applied alloys for orthopaedic applications is very important for achieving improved biocompatibility. It is important that biomaterials will not introduce fibrous tissue barriers during contact with healthy bone. This permits the bone to grow closer to the surface of the implant and to fill in grooves or pores that have been deliberately introduced to a surface according to the Wolf law of bone growth to enable the device to become more firmly embedded.General requirements for orthopedic implantsArtificial solutions for human body disabilities must be based on the mechanical design and actually available materials, which meet special requirements. These are mechanical, biological and chemical requirements that must be fulfilled for them to be used for the human body in line with the principles of medicine describing the particular features of biomaterials. The above review of the literature reveals complex issues, which are related with medical, mechanical and material problems in hip implants. Biomaterials used for hip prostheses should exhibit the following properties [6], [7], [9], [10], [13], [21]:High biocompatibility (nontoxic and non-carcinogenic, chemically stable and corrosion resistant);Nonmagnetic;The surface femoral stem should be bioactive in terms of osteointegration;Adequate mechanical properties (Young modulus, yield strength, tensile strength, ductility, hardness, toughness, fatigue strength and abrasive wear);The femoral stem should be able to endure large and variable stresses in the human body environment relating to a large number of stress cycles, i.e. fatigue (e.g. during 10 years, assuming that the distance of 10 km in a day yields about 35×106 mechanical cycles, i.e. steps) with additional stress corrosion, stress-enhanced diffusion and even stress-induced phase transformation;Surface layer with gradient-like modulus of elasticity, i.e. with gradient-like porous distribution in surface layer;The surface of implants should be free of imperfection, such as tool marks, scratches, nicks, cracks, cavities, burrs and other defects;Surface texture as surface quality factor and surface geometry classification may vary from very smooth for bearings to rough for the femoral stem and socket. All this depends on the design, material, application and the area involved in the metallic implant (e.g. the fatigue crack initiation is significantly lowered when the surface is highly polished, whereas a rough surface is demanded for better osteointegration);Well-balanced chemistry and proper metallurgical conditions of the alloy, i.e. proper chemical, phase and microstructure composition, absence of segregation, nonmetallic inclusions and martensitic phase;Availability, with reasonable balance between quality, durability and price.Mechanical and materials requirements for orthopedic implantsOne of the goals in the development of new materials for hip implants was and still is an optimal ratio of the elasticity modulus of metallic, ceramic and polymer parts and the bone. The modulus is an important concern in the orthopaedic application of biomaterials in terms of the mechanical interaction between the implant and the bone. The stiffness of the material can be described by Young’s modulus of elasticity (E). The ratio of stress (s) to strain (e) in the elastic deformation region is known as the modulus of elasticity (E), or Young’s modulus. Young’s modulus is the tangent of slope angle of the σ=Ee strain line function in the elastic region of the tensile test. The constant of the proportion between the stress and strain in the linear elastic region during tension is known as Young’s modulus and defined as(1)Modulus of elasticity E=σx/εx,σx= Eεx(here E as stiffness),  εx= σx/E and εy= −νσx/E (here 1/E is susceptibility)$$\matrix {{{\text{Modulus of elasticity }}E = {\sigma _x}/{\varepsilon _x},} \hfill \\ \matrix {\sigma _x} = {\text{ }}E{\varepsilon _x}({\text{here }}E{\text{ as stiffness}}),{\text{ }}{\varepsilon _x} = {\text{ }}{\sigma _x}/E{\text{ and }} \hfill \cr {\varepsilon _y} = {\text{ }} - \nu {\sigma _x}/E{\text{ (here }}1/E{\text{ is susceptibility)}} \hfill \cr \hfill \\ } $$The modulus of elasticity (E) is essentially a measure of the material stiffness and is determined from the tensile test or by the elastic internal dumping method (internal friction) and by ultrasonic wave velocity measurements.The resistance to fatigue crack nucleation depends on stress, microstructure and modulus of elasticity according to the Paris law and Irvin and Weibull considerations [22], [23]. Fatigue cracks usually nucleate at the surface of the stressed metal, where the stresses are at a maximum taking into account surface residual stresses and the principle of superposition of any existing stress component. In contact loading, the Hertz and Bielajew point with the maximum reduced stress must be taken into account. Any design or manufacturing defects and tensile surfaces with residual stresses encourage nucleation and development of fatigue cracks. The fatigue resistance of the material can be expressed as the crack growth rate (da/dN) related to applied stress, strength (s, K) and Young modulus (E) of the material [23]:(2)dadN=CΔKm=Ca(KE)3.6$$\frac{{da}}{{dN}} = C\Delta {K^m} = {C_a}{\left( {\frac{K}{E}} \right)^{3.6}}$$whereda/dN is the crack growth rate [mm/cycle]C, Ca and m are the experimental constantsΔK=Δσ is the stress intensity factor during fatigue factorwhere sharp crack occurs, the existence of a critical stress for rapid crack propagation is given by the relation(3)σcrict=Eγ/πa$${\sigma _{{\text{crict}}}} = \sqrt {E\gamma /\pi a} $$where σcrit is the critical tensile stressE is Young’s modulusγ is the energy per unit area of new surfacesa is the crack length for a surface crack and 2a=for an internal crackThe stress from the applied loading must be lower than the tensile strength of the material (Table 1). If the tensile strength of the material increases, the resistance to fatigue also increases by the crack growth rate decrease [Eq. (2)]. If Young’s modulus is larger, the crack growth rate is smaller and the critical stress to develop crack must be larger [Eq. (3)].Table 1:Young modulus and mechanical properties of the bone and selected biomaterials.Properties/materialYoung’s modulus E, GPaPoisson’s ratioTensile ultimate strength Rm, MPaTensile yield strength Re, MPaRe/RmBending strength Rg, MPaUltrahigh molecular weight polyethylene (UHMWPE)55–1700.42>30>20––Bone cement3.80.460550.9–Compact bone30–120––160Ti CP Grade 4/annealed1050.37560–880380–5200.7340Ti-6Al-4V/annealed1150.41>890>7300.9650Cobalt alloys (Co-28Cr-6Mo)/annealed2300.36>600>5000.6–Bio-steels (see Chap. austenitic steels)2100.29480–900180–6000.6250Reciprocal relation appears when there are differences in an elastic module of the bone and the implant. Significant differences in the elastic moduli of the bone and the implant may lead to stress concentrations at their interface, i.e. the stress threshold. An insufficient load transfer from the artificial implant to the adjacent natural bone may result in bone desorption and loss of the bone-implant joint, and finally, the whole hip implant can be rejected. The low modulus of alloys reduces the mismatch between the modulus of the implant and that of the bone, which in turn reduces the so-called “stress shielding phenomenon”, which can be responsible for potentially damaging resorption of bone on the inner surface of the natural femur. Stress shielding is related to the difference in flexibility/susceptibility or stiffness between natural bone and the implant alloy. The lower modulus of titanium alloys is a positive factor in reducing bone resorption. However, the modulus of elasticity should not be too low. An implant made from alloy with the too low Young’s modulus of elasticity will have low resistance to nucleation of fatigue cracks and larger cracking velocity [Eq. (3)]. From this point of view the modulus of elasticity (E) should be in the optimum range owing to the selection of proper materials and/or proper production technology.Recent attempts to minimize the modulus of elasticity and improve biocompatibility have led to the introduction of some metastable β titanium alloys with elastic modulus values as low as about 70–80 GPa [24], [25], [26]. However, there has been, and still is, concern about too large elastic modulus of implant alloys compared to that of compact bone (30 GPa). The high difference in the modulus of elasticity of the alloy and of the bone causes stress incompatibility by stress, stress concentration and stress threshold. The modulus of elasticity should be optimum which assures better transfer of stresses between femoral stem with flange and a porous coating of alloy with gradient porosity, which means a gradient-like distribution of Young’s modulus. The modulus of elasticity depends on the degree of porosity. The modulus of elasticity decreases with increasing the porosity of coating from the implant surface to the bone.Amongst metallic biomaterials, the elastic modulus of Ti-6Al-4V alloy (E=110 GPa) is much lower than that of 316L (X3CrNiMnMo18-14-5-3) stainless steel (E=210 GPa) or Co-Cr-Mo alloy (E=230 GPa). Other mechanical properties like hardness, ultimate tensile strength (UTS), Rm, yield strength (R0.2), bending strength (Rb), torsion strength, toughness and fatigue strength of the alloys considered should be also measured and analyzed (Table 1).Mechanical properties depend on the microstructure which is characterized by grain size, grain crystallographic orientation, crystal lattice defect density and phase composition. The last one is formed by phase transformation due to heat, thermochemical and thermomechanical treatments. Metal forming, i.e. cold and/or heat plastic deformation, are important technological operations which influence lattice defects, grain size and their orientation. Segregation of alloying elements can result in the formation of non-uniform phases and microstructure composition with different crystallographic structures, lowered corrosion resistance and fatigue strength and increased residual stresses and pitting corrosion. Agglomeration or large amounts of other phases can reduce the fatigue life of implants. Segregation of non-uniform clusters acts as stress threshold in the microstructure. An alloy with segregated alloying elements is sensitive to nucleation of fatigue cracks. The fatigue or stress corrosion damage and abrasive wear are the major causes for failure of orthopaedic implants. The fatigue strength yield point ratio of applied biomaterials is small and ranges between 0.3 and 0.7 with additional large safety coefficient which makes the calculations of mechanical strength and total reduced stress rather complex and difficult.Biomedical, mechanical and material engineering of hip implantsThe natural constitution of the hip joint, loading stresses, residual stresses and corrosion, mechanical design of human bearing and the materials which are available define problems to solve for artificial hip implants. Mechanical, chemical and physical properties as well as biocompatibility requirements define suitable solutions for this complex problem. Typical components of hip implants are shown in Figure 1.As shown in Figure 1, a typical hip prosthesis consists of a femoral stem, a femoral head and socket. The femoral stem is usually manufactured from Co-Cr-Mo or Co-Ni-Cr-Mo alloys or from titanium base alloys (Ti-6Al-4V or new one Ti-Nb-Zr). The head is made of highly polished Co-Cr-Mo alloy, of Al2O3 sintered ceramic composite or of stabilized ZrO2 with Al2O3 and yttrium oxide additives [10]. The metal-backed shell and the screws are made of titanium alloy Ti-6Al-4V or CP titanium (Tables 1 and 2). The socket (cup) is made of UHMWPE usually with metallic shell. Modular designs, where the stem and head are made of two materials, are common [11]. However, when two different metallic alloys are applied, problems with galvanic corrosion should be considered and taken into account.Table 2:Chemical composition of commercially pure titanium grades in wt.% according to ASTM F 67.DesignationTiC (max)Fe (max)N (max)O (max)H (max)OtherASTM Grade 199.50.080.200.030.180.15–ASTM Grade 299.20.080.500.030.200.15–ASTM Grade 399.10.080.250.050.300.15–ASTM Grade 499.00.080.500.050.400.15–ASTM Grade 799.20.100.300.030.250.150.2 PdProduction technologyMechanical and biomedical properties depend on chemical, phase and microstructure composition. The technological processes which are applied influence characteristics of the above materials. Therefore, knowledge about the relationship between particular technological operations and the properties considered is an important key to understanding the technology of production of implant parts and their expected features. Generally, metallic parts are produced by casting, metal forming (forging, rolling and pressing), powder metallurgy (compacting and sintering) and various types of heat treatments and machining. The most recent manufacturing method, i.e. “3D printing” as an additive technology, is under development. Commonly, metallic parts of hip implants are manufactured by forging. Investment casting of hip stem is used as manufacturing method which is an alternative to forging. The main disadvantages of parts produced by investment cast are lower mechanical properties than those of forged parts. Forged hip stems have a higher yield strength, tensile strength and fatigue strength than those made by a cast. This is due to the proper microstructure of alloy resulting from large plastic deformation during forging. The microstructure of cast alloys, called primary microstructure, is generally worse due to non-uniformity like segregation of alloying elements and different grains size in different places. Mechanical properties of forged components are better than those produced by casting or 3D printing technology. The femoral stem is usually forged as α+β titanium phases at a temperature of 955°C. Forged stem parts are then annealed at recrystallization temperature. The microstructure is either mill annealed or fully equiaxed depending on the parameters of annealing according to the principle: the higher the temperature of annealing, the larger the grain size and the smaller the mechanical properties. Mechanical properties of this alloy can be controlled over significant range through hot and/or cold working and heat treatment.The compacting and sintering technology is used to produce ceramic components. The following technological operations are heat or thermochemical treatments including stress relaxation, annealing, super saturation, ageing, carburizing and nitriding. Grinding and polishing as the final machining are applied to obtain the final shape and quality of a surface. The introduction of compressive stresses to the surface layer by thermochemical treatment (for example, by nitriding or carbonitriding) or shot penning increases resistance to fatigue cracking. The majority of all the above considerations relate to metallic components of hip implants; however, the contemporary modular structure of hip implants still needs to be developed in the field of design, manufacturing technology and materials. The advanced metallic, ceramic, composite and polymer materials are an important key to the progress of hip prosthesis and of biomedical engineering in general.Introduction of titanium and titanium alloys for biomedical engineeringThe selection of titanium and titanium alloys for hip implants is determined by a combination of most favourable characteristics including high strength to weight ratio, outstanding resistance to corrosion by body fluids, high biocompatibility, low density and the capacity for joining with bone, i.e. osteointegration. Titanium alloys have outstanding resistance to corrosion by body fluids, which is superior to that of stainless steels. Although titanium is highly reactive to oxygen, the creation of passive oxygen layers and solid interstitial solution exhibits excellent corrosion resistance in the human body environment. Sliding contact should be avoided without modifying the surface by thermochemical diffusion treatment (such as nitriding see, e.g. Figure 2).Figure 2:X-ray diffraction pattern of nitrided Ti-6Al-4V alloy with Bragg-Brentana and grazing angle (α) diffraction geometry for comparison of phase composition versus X-ray depth of penetration (z), (for BB z=5–18 μm; for α=2, z=1.1 μm; for α=4, z=2.4 μm; and for α=9, z=4.8 μm) [27].Titanium exists in two allotropic crystallographic forms: Tiα, which has the hexagonal closed packed crystallographic structure, and Tiβ form, which has the body cubic (BCC) structure. In pure titanium, the α phase is stable up to 882°C. Above 882°C the α→β transit temperature appears. The hexagonal α phase of pure Ti is transformed in heating to the BCC β phase at a constant temperature of 882°C. This phase transformation temperature depends on the chemical composition of titanium alloys [28]. Commercially pure titanium which is unalloyed ranges in purity from 99.5 to 99.0% Ti (Table 1). The main elements in unalloyed titanium are iron and interstitial elements like carbon, oxygen, nitrogen and hydrogen [2], [4], [19], [21]. The chemical compositions of the principal grades of titanium are listed in Table 1. Commercially pure titanium can be considered as an α phase alloy in which the oxygen content determines the grade and strength. Titanium sponge due to metallurgical technology contains oxygen at a certain level, but its amount is adjusted to modify the strength of commercially pure titanium. In this case, oxygen can be treated as an “alloying element”. Carbon, nitrogen and hydrogen are present as metallurgical impurities in titanium. Oxygen within solution limits can be used as an interstitial strengthening element. Hydrogen is readily absorbed by titanium alloys and is detrimental because it causes so-called hydrogen embrittlement by brittle needle-like precipitation of titanium hydrides (TiH2). Oxygen, nitrogen and hydrogen, on the other hand, can improve biocompatibility. Nitrogen, carbon, oxygen, boron and hydrogen form interstitial solid solutions because of the large difference between the atom size of titanium and these elements. There is a reasonable difference in solubility of these elements in α and β titanium [28]. Titanium and its alloys react with several interstitial elements including gaseous oxygen and nitrogen, and such reactions occur at a temperature well below the respective melting points. This chemical affinity is used in the thermochemical treatment of titanium alloys.Commercially pure titanium of ASTM grades have low strength but higher corrosion resistance than titanium alloys. The addition of 0.2% Pd to commercially pure titanium improves its strength and corrosion resistance in active reducing media and is designated as Grade 7 (Table 2).Although interstitial elements increase the strength of titanium, they are detrimental to toughness as measured by the notch impact test. Therefore, when high toughness is desired for certain applications, the alloy should be produced with extra low interstitials (ELI). These alloys are referred to as ELI alloys [27]. Strain hardening of titanium alloys is strongly sensitive to the rate of deformation. Increasing rate of deformation increases significantly the strain hardening rate. In order to obtain a large deformation of titanium alloys, the process is carried out in hydraulic presses with a slow deformation rate. A slower deformation rate produces larger deformation of titanium alloys without cracking. Plastic deformation is realized with a dislocation slide mechanism suitable to the crystallographic lattice of both α and β phases [29]. The twining system of deformation appears in both titanium phases as well.The titanium α/β alloys are most often used in the annealed condition. Their microstructure and mechanical properties may differ depending on whether or not prior plastic deformation was carried out above or below the α/β phase transformation temperature. A typical forging temperature of α/β Ti-6Al-4V is 955°C. An inert gas, e.g. argon can be introduced into one part of the mould cavity. During the heat treatment, the high vacuum atmosphere is used, 10–5 torr). Above 535°C, titanium absorbs oxygen and forms oxide compounds (TiO, Ti2O3 and TiO2) and a solid solution as a subsurface layer. All of them can improve mechanical properties from one side and nucleate surface and brittle cracks from another side. Titanium also absorbs hydrogen, and it is necessary to ensure that furnace atmospheres are hydrogen free during melting, heat and thermochemical surface treatments. Titanium alloys are particularly susceptible to galling, i.e. wear due to friction, during hot or cold working, which causes surface damage. Therefore, proper lubricants should be applied during thermomechanical cold or hot working. During these processes, graphite or molybdenum disulphide is used, whereas glass powder may be used for a more severe process like hot extrusion [27].Physical metallurgy and surface layers of biocompatible titanium alloys for hip implantsTitanium alloys are sensitive to contact in friction couple due to sticking or galling, which may lead to fretting. For this reason, titanium alloys have low resistance to fretting. Fretting occurs when two surfaces are in contact and small amplitude relative oscillation motion is present. Fretting damage in titanium alloys can significantly reduce fatigue life. Fretting is a major concern for hip implants in joints between femoral stems and heads.Fretting corrosion is a serious problem in implants because it is the main mechanism of gradual release of metal ion to different tissues in the body. Fretting corrosion and abrasive wear cause pain and excessive accumulation of wear particles which results from prosthesis losses. The extensive release of metal ions from hip prosthesis can result in adverse biological reactions [12], [19].Two primarily problems arise when attempting to coat titanium alloy with TiN, TiC and TiCN layers. The thermal expansion coefficients of the coating and of the titanium substrate are different, and as a result, large thermal stresses are generated at the interface during the coating deposition process [30]. Such coating also has a different microstructure, crystallographic structure, different mechanical properties (strength, hardness, modulus of elasticity etc). The coating layer of the femoral head should not be too hard. Under heavy load and residual stress, the coated layer will be cracking. These stresses can lead to cracks at the interface coating/substrate. In addition, chemical reactions between the coating and titanium alloy can weaken the metal in the vicinity of the interface, reducing strength and corrosion resistance of the coating system. Most advanced coatings were elaborated with the application of hydroxyapatite and some other ceramic compounds [19], [31].Therefore, diffusion thermochemical treatment is a better solution in order to produce hard, highly resistant to wear and gradient like mechanical properties versus depth from the surface. Here is an example of a thermochemical ion nitriding of Ti-6Al-4V alloy (Figure 2). Very clean nitrogen (N2) was introduced to vacuum furnace chamber. Under high voltage, electrical discharge occurs and the nitrogen is dissociated, ionized in the form of glow discharge which produces single ions (in statu nascendi) according to the following reaction:N2→N2++N2++4e${{\text{N}}_2} \to {{\text{N}}^{2 + }} + {{\text{N}}^{2 + }} + 4{\text{e}}$Only an active nitrogen in “statu nascendi” can diffuse to the metal surface layer of metal. The diffusion ion nitriding treatment (plasma nitriding) was carried out in controlled atmosphere (containing a mixture of nitrogen ions and argon) at a temperature of 700°C. The phase composition varies versus thickness of the surface layer, and some level of residual stresses was measured (about −522 MPa) with GID-sin2ψ method [30]. The qualitative X-ray phase analysis showed the dominating phase component as Tiα, TiN and Ti2N (Figure 2). The two latest nitrides appear in a thin surface layer of a nitrided Ti-6Al-4V alloy with the plasma thermochemical process. The investigated surface layer was exposed with X-ray grazing angle (α) diffraction geometry with incident beam angle α=2 and α=4°. The applied X-ray diffraction geometries are destined for thin surface layer analysis with different thicknesses. This means that very thin surface layer thickness of 2 and 4 μm consists of TiN and Ti2N, and deeper layers as a substrate under them consist of hexagonal phase Tiα(N) as a solid solution with a gradient-like nitrogen distribution (Figure 2). Similar experiments and examinations carried out on CP4 titanium [18] under a little bit different conditions confirmed phase composition and compressive residual stresses.Classification of titanium alloysTitanium alloys are classified into categories according to the microstructure and phase composition (structure) retained at room temperature after solidification and other metallurgical processes. Both characterizations depend on heat and various kinds of thermomechanical treatments. According to the phase composition (as structure composition), titanium alloys can be divided intoα alloysnear α alloysα+β alloysβ alloysnear β alloysmetastable β alloysThe most known Ti-6Al-4V titanium alloy contains 6 wt.% aluminium and 4 wt.% vanadium. It was invented in 1946 in the USA as a most advanced alloy for aeroplane and cosmos industry applications. Soon afterwards, this alloy was experimentally introduced as a good biocompatible alloy for medical application. Gradually, aluminium and later vanadium were substituted by elements more friendly for the human body (Table 3). A new generation of titanium alloys with niobium was introduced in 1979 [5], [25], [26], [32], [33], [34]. New β or near β titanium alloys contain biocompatible alloying elements: Nb, Zr and Ta. They have a lower modulus of elasticity as compared to α+β alloys. The β alloys have excellent formability, cold rolling capabilities, good ductility and formability. The main disadvantage of β alloys in comparison with α+β alloys is their higher density. Most recently developed alloys based on the Ti-Nb-Zr-Ta alloying elements appeared as β alloy has the lowest modulus of elasticity, E=70 GPa [3], [24].Table 3:Selected experimental and applied titanium alloys for implant applications.Alloy designation and typeDeveloped inα+β alloysTi-6Al-4VUSATi-6Al-4VELIUSATi-6Al-7NbSwitzerlandTi-5Al-2.5FeGermanyTi-3Al-2.5VUSATi-15Zr-4Nb-2Ta-0.2PdJapanTi-5Al-3Mo-4ZrJapanTi-15Sn-4Nb-2TaJapanβ alloysTi-15Mo-5Zr-3AlJapanTi-29Nb-13Ta-4.5ZrJapanTi-13Nb-13ZraUSATi-12Mo-6Zr-2FeUSATi-16Nb-10HfUSATi-35Nb-7Zr-5TaUSATi-15Mo-2.8Nb-0.2Si-0.26OUSAaThe only alloy with homogeneous microstructure is single-phase β. Alloy with segregation of Nb and Zr is not acceptable. Segregation of alloying elements (Nb and Zr) causes not homogeneous microstructure sensitive to corrosion and fracture.Alloy Ti-13Nb-13Zr was originally developed for medical implant applications [24]. This alloy combines a high biocompatibility, high strength, lower elastic modulus than alloy Ti-6Al-4V, excellent hot and cold workability and superior corrosion resistance. The Ti-13Nb-13Zr alloy is a metastable beta titanium alloy developed for use in biomedical implants that combines a high biocompatibility, superior corrosion resistance, high strength, excellent hot and cold workability and lower elastic modulus. Mechanical properties of this alloy can be controlled over significant range through hot working, cold working and heat treatment. The mechanical properties of Ti-13Nb-13Zr can vary by thermomechanical treatment. This alloy contains a high percentage of high-density alloying elements. Because of the high density of zirconium and niobium, there is a problem of segregation during melting and solidification of alloys containing these elements.Also, because of the high density of niobium and zirconium, special techniques of mixing liquid metal during melting and casting process are required to prevent segregation of alloying elements. Niobium and zirconium have larger atomic diameter than titanium. Special thermochemical treatment which provides point defects and dislocations should be applied to increase the diffusion rate of alloying elements in order to obtain the proper response during homogenization treatment. Segregation of alloying elements cannot be tolerated in metallic implants. Segregation of alloying elements decreases resistance to corrosion and fatigue crack initiation, lowers resistance to an impact load and fracture resistance and decreases biocompatibility. The newest development of titanium alloys is directed to Ti-Au composition with an expectation of better mechanical and biomedical properties. Another titanium base alloy with nickel, nitinol, is also an important biomaterial [35], [36].Cobalt base alloys for hip implantsCobalt-chromium alloys have good resistance to pitting and crevice corrosion in the human body. The main cobalt alloys (Table 4) used for hip implants areASTM F75, (Co-28Cr-6Mo) casting alloyASTM F90, (Co-20Cr-15W-10Ni) wrought alloyASTM F999, (Co-28Cr-6Mo) thermomechanically processedASTM F562, Co-35Ni-20Cr-10Mo wrought alloyASTM F 1058 (Co-20Cr-15, 5 Ni-7Mo-Fe) wrought alloyMP35N, (35Co-35Ni-20Cr-10Mo) thermomechanically processedTable 4:Chemical composition in wt.% of selected cobalt base alloys used for hip implants (Co balance).ASTM designationUNS no.CrMoNiFeCSiMnWPSOtherF75R3007527.0–30.05.0–7.01.00.750.351.01.00.200.0200.010.25 N, 0.30 AlF90R306519.0–21.0–9.0–113.0 max0.05–0.150.400.4014.0–16.00.0400.03–F562R3056319.0–21.09.0–10.533.0–37.01.0 max0.025 max0.150.15–0.0150.011.0 TiF563R3056318.0–22.03.0–4.015.0–25.04.0–6.00.050.501.03.0–4.0–0.010.50–3.50 TiF799R3153726.0–30.05.0–7.01.00.750.351.01.0–––0.25 NBiophysics and physical metallurgy of cobalt alloysCobalt is beneficial for humans because it is a component of vitamin B12 (cobalamin), which is essential for human health. However, too high concentrations of cobalt may be toxic to health. A cobalt-based alloy was first used as a material for hip implants in the USA in 1936. At temperatures below 417°C cobalt exhibits a hexagonal close-packed structure. Between 417°C and its melting point of 1493°C, cobalt has a face-centred cubic (FCC) structure. The melting point of cobalt is 1768°C. The elastic modulus of cobalt is about 210 GPa in tension and about 183 GPa in compression. Cobalt has a density of 8.90 g/cm3. Typically, the microstructure of cobalt-based alloys consists of FCC as (γ) phase as matrix and various types of carbides. A strengthening in cobalt-based alloys is obtained primarily through a combination of solid solution strengthening (by Cr, Mo and Fe atoms) and carbide precipitation. Strength and toughness depend on the amount of solved foreign atoms and amount, size and distributions of carbides. A fine dispersion of carbides contributes significantly to the strength of these alloys. Dislocation movement is strongly impeded by carbide precipitations. Carbide precipitations are barriers for dislocation movements. This carbon content in these alloys has a significant effect on the number of carbides. In general, three main types of carbides can be present in the microstructure of cobalt-based alloys [8], [14], [15], [16], [17]:M23C6 carbides, where “M” is mostly chromium but can be substituted by tungsten and molybdenum;MC carbides, where M stands for the reactive metals tantalum, titanium, zirconium and niobium;M6C carbides, where “M” stands for tungsten or molybdenum; these carbides form when the tungsten or molybdenum content exceeds about 5 at.%.Components made of cobalt alloys are designed as perfect spherical head of hip bearing so that the metal always articulates against a low friction plastic (e.g. UHMWPE or high cross-linking polyethylene), which provides smooth relative movement, small friction and minimum abrasive wear. The recent reports on cobalt alloys applied for hip implants show an easy diffusion of Co ions from the hip implant to body fluids. This type of implementation with cobalt alloy implants caused excessive concentration of Co in blood [20]. Therefore, some limitations in this type of application have been prepared.Austenitic stainless steelsThe main components of stainless steels are iron, nickel and chromium. Generally, stainless steels are divided into three groups: ferritic, martensitic (tool steels) and austenitic. Only the last type of stainless steels called bio-steels is applied for a hip prosthesis for short-term use. Austenitic stainless steels were first used as a material for hip implants in the USA in 1926.Nevertheless, iron ions are important components of blood; other chemical components are rather harmful to human body.Austenitic stainless steels are used for implant applications because they are relatively inexpensive. Passivation of stainless steel implants is enhanced by nitric acid. Austenitic stainless steels are not sufficiently corrosion resistant for long-term use as an implant material. Pitting corrosion in austenitic stainless steels is usually caused by chlorine ions and starts at nonmetallic inclusions (sulphur inclusions) and other heterogenic places. Molybdenum is added to stainless steels usually in amounts of 2–3% to increase resistance to pitting corrosion in saline environments. Nitrogen-strengthened alloys are used for the bone plate and bone screws.Austenitic stainless steels have excellent formability, and their response to deformation depends on the nickel content. Formability increases with an increase in the content of nickel. Magnetic alloys should not be used in the human body because they could become dislodged in the strong magnetic field of magnetic resonance imaging. Pitting corrosion in the stainless steels occurs when the protective film is broken, exposing the steel to Cl− ions. Austenitic stainless steels for hip implants can be divided intolow carbon stainless steelsnitrogen-strengthened stainless steelsnickel-free stainless steelsExamples of major ASTM austenitic stainless steels for implant materials:Low carbon stainless steels: ASTM F138/139 316 L, (21Cr-10Ni-3Mn-2.5 Mo). This steel contains 0.25–0.50% N.Nitrogen-strengthened stainless steels: ASTM 1314 (22Cr-12, 5Ni-5Mn-2.5Mo-0.30 N). This steel contains 0.20–0.40% N.Nickel free stainless steel and nitrogen-strengthened steels: ASTM 1314 F2229 (23Mn-21 Cr-1Mo –0.95N).The substitution of nickel atoms is important due to its harmful activity in the human body. Because of that, Ni-free stainless steels were elaborated (Table 5).Table 5:Chemical compositions (in wt.%) of austenitic stainless steels used for surgical implants according to the ASTM specifications (Fe balance).aASTM designationCCrNiMnMoCuNPSOtherF1380.0317.00–19.0013.0–15.002.002.25–3.000.500.100.0250.010–F13140.0320.50–23.5011.50–13.504.00–6.002.00–3.000.500.20–0.400.0250.0100.10–0.30 Nb; 0.10–0.30 VF15860.0819.50–22.009.00–11.002.00–4.252.00–3.000.250.25–0.500.250.0100.25–0.80 NbF22290.0819.00–23.000.1021.00–24.000.50–1.500.250.90 min0.030.0100.95 NaSingle values are maximum values unless otherwise indicated.SummaryThe review, as presented above, i.e. titanium and cobalt alloys, bio-steels, their specific treatments, surface layers and coatings on alloys used for hip prosthetic can be summarized as follows:The synergy of some branches of science and technology, i.e. medicine (surgery), mechanical design and material science and technology has led to progress and achievements in hip implants.All of the chemical elements of the materials applied in implants are building components of the human body, but some of them are even poisonous and others, like aluminium, vanadium and nickel, are classified as toxic elements.Biomedical implant reliability depends largely on the corrosion, abrasive wear and fatigue strength/resistance of the materials used in production technology.The capacity for joining implant with the bone, i.e. osteointegration, which is direct bone anchorage to an implant into bone with no fibrous capsule interface, makes it possible to transmit the loading forces directly to the bone.All the beneficial passive layers on metals can be mechanically damaged by the fretting of metal against metal.Titanium alloys are used for hip implants due to excellent combination of mechanical properties, low density, acceptable tissue tolerance, high strength to weight ratio, outstanding resistance to corrosion by body fluids, high biocompatibility, low density and nonmagnetic properties.Titanium induces the formation of a fibrous tissue barrier when placed in contact with a healthy bone. This permits the bone to grow closely on the surface of the femoral stem and fill pores that have been deliberately introduced to facilitate bone ingrowths.Contaminations of titanium alloys by interstitial elements such as hydrogen and oxygen during fabrication (melting, thermomechanical treatment and thermochemical surface hardening) must be avoided because these elements have an embrittling effect on titanium.In the presence of hydrogen, titanium hydrides (TiH2) may form into long thin needles in titanium alloys and these hydrides decrease resistance to fracture.Austenitic stainless steel, particularly nickel-free stainless steel with nitrogen strengthened like ASTM 1314 F2229 (23Mn-21Cr-1Mo-0.95N), can be used for some elements of hip implants.Cobalt-chromium alloys are the most commonly used metals for current metal on polymer implants. A typical microstructure of cobalt base alloys consists of FCC (γ) phase matrix and carbides of various types.Hydroxyapatite is widely used as bioactive porous coating of the femoral stem and the metallic shell in hip implants.A special type of polyethylene, i.e. special grade UHMWPE with additional cross linking bonding, is widely used (as socket) in hip implants as the load-bearing material (in hip implants with combination metal on polymer).The final decision about the type and material of hip implants made in the hospital by medics is supported by many additional factors like accessibility, availability of the material, the patient’s weight and age, etc.ConclusionsThe synergy of medicine, biochemistry, biophysics, mechanical engineering, materials science and informatics succeeds in achieving real progress in orthopaedic surgery of hip joint implants.Although the majority of chemical elements of the materials applied to make implants are building components of the human body, it is usually harmful if their concentrations in some tissues, bone or in body liquids are over the natural level.The fatigue and stress corrosion damage are the major causes for the failure of orthopaedic implants. The fatigue strength and yield point ratio of biomaterials applied are small and range between 0.3 and 0.7, which, with additional large safety coefficient, cause mechanical strength and total reduced stress calculations to be rather complex and difficult.The majority of parts of hip joint implants are made of metallic materials due to heavy and cyclic loaded bearings. Additionally, they work in the bioactive environment. Titanium alloys are most advanced materials in this type of application; however, Co-Cr-Mo alloys and austenitic stainless like bio-steels mated with appropriate metallic alloys are the main materials for components of hip implants.The natural constitution of the hip joint; the stresses resulting from loading, residual stresses and corrosion; mechanical design of human bearing; and accessible materials define problems to solve for artificial hip implants.The contemporary modular structure of hip implants still needs to be developed in the field of design, manufacturing technology and materials. The advanced metallic, ceramic, composite and polymer materials are an important key to the progress of hip prosthesis and of biomedical engineering.NoteThe information in this article is for informational and educational purposes and is not meant as medical advice or recommendation. Only a qualified orthopaedic surgeon can determine which material implant system is best for an individual person. There are many factors that the surgeon uses when recommending hip implants (state of health, weight, age, life activity, anatomy etc.). Any questions and concerns regarding specific type of materials and implants should be discussed with the professional orthopaedic surgeon. Every patient’s case is unique, and each patient should follow specific instructions of his or her doctor. The information in this article does not replace orthopaedic doctor’s specific instructions.Acknowledgments:The authors would like to thank MSc Kamil Nawojowski for his support.Author contributions: All the authors have accepted responsibility for the entire content of this submitted manuscript and approved submission.Research funding: The results of research and this paper were supported by the Faculty of Metals Engineering and Industrial Computer Science, Department of Physical and Powder Metallurgy of AGH-University of Science and Technology in Krakow under project no 11.11.110.299.Employment or leadership: None declared.Honorarium: None declared.Competing interests: The funding organization(s) played no role in the study design; in the collection, analysis, and interpretation of data; in the writing of the report; or in the decision to submit the report for publication.References1.Niinomi M. Recent titanium R&D for biomedical applications in Japan. 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