Introduction

Biomaterials, in general terms, are defined as natural or synthetic materials that can be used to augment or replace the biological structure in partial or full capacity to improve the life quality and longevity of an individual.1 Early applications of biomaterials can be dated back to ancient civilizations. Historians suggest that the early Romans used to apply glass to recover eye injuries, Chinese and Aztecs utilized the noble nature of gold in dentistry, and Egyptians and Indians used linen as a suture to seal wounds.2 Eventually, the growing clinical need and the collaborative effort of material scientists, biologists and chemists led to a tremendous improvement in the material and application range of such materials. Nowadays, biomaterials are used in a variety of body parts. A brief overview and description are shown in Fig. 1.

Fig. 1
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Hard and soft tissue biomaterials applied to the human body.3,4,5,6,7,8,9,10,11

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Given the complexity of the human body, it is always a challenge to fulfill all the necessary requirements for a foreign material to perform as well as the native organs. Just to provide an overview about the challenges, it should be noted that the pH varies from 1 to 9 in body fluids of various tissues.12 Additionally, the mechanics of the body helps to generate massive force concentrations in several body parts, which can vary to a large extent, depending on the individual. In day-to-day life, bones are subjected to maximum stress up to ~ 4 MPa, whereas tendons and ligament stresses are around 40 MPa to 80 MPa. The mean load subjected to hip joints can be up to ~ 3 kN, which increases up to tenfold during jumping. Another crucial point is the fluctuation and repetition of these stresses depending on the activity such as sitting, standing, stretching, climbing, jogging and running.13 Even if a material can fulfill the structural compatibility desired, it still must also be biocompatible to interact with the human tissue without any negative effects such as inflammation or creation of undesirable chemical components. Biocompatibility can be defined as the ability of the material to show favorable response in a given biological environment without having any risk of rejection by the immune system, toxicity, injury or any other undesirable effect.14,15,16

As it has not yet been possible for a single material to fulfill all the above-mentioned requirements, the biomaterials are classified based on the structural and biocompatibility to allocate them in applications according to their properties. The structural compatibility difference divides biomaterials into (1) natural materials, (2) metals, (3) ceramics, (4) polymers and (5) composites, which are developed by different combinations of the above.17,18 Whereas, the biocompatibility difference divides biomaterials into (1) bioinert [e.g., titanium (Ti) and its alloys, some steel grades, alumina (Al2O3), titanium nitride (TixNy), titanium dioxide (TiO2), zirconia (ZrO2)], (2) bioactive [e.g., hydroxyapatite (HAp), Ca10(PO4)6(OH)2, bioactive glasses (e.g., CaO-SiO2-P2O5-Na2O) and glass ceramics]19,20,21 and (3) bioresorbable [e.g., tricalcium phosphate (TCP), Ca3(PO4)2, some alloys of magnesium (Mg), zinc (Zn) and iron (Fe)].19,22 The bioinert materials are the least biocompatible, which show minimum interactions with surrounding tissues after being placed inside the body. Bioactive and bioresorbable materials interact with surrounding tissues to a high extent with no harmful effects, where bioresorbable materials even get resorbed by the hydrolytic breakdown in the body and slowly get replaced by the advancing tissue (e.g., bone). The chemical by-products during this process are absorbed and released by the body via various metabolic processes.23,24 The challenge—among others—is to adapt the corrosion of the implanted material to the growth of the tissue.

Biomaterials are primarily used to replace hard and soft tissues that have been damaged or destroyed.25 As the mechanical properties required for hard tissues are quite high compared to soft tissues, the material selection for such applications also varies to a high extent.12,26 In Table I, some mechanical properties of hard tissues and different classes of biomaterials along with their respective biocompatibility for hard tissue applications are listed.

Table I Classification of biomaterials based on structural and biocompatibility
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The biocompatibility of natural biomaterials is evidently the highest, but their lack of mechanical strength renders them unsuitable for hard tissue applications.17,80 The high elastic modulus and tensile strength provide metals and alloys a better load-bearing capacity and thus are being widely used for hard tissue applications.81,82 However, their bioinertness (excluding Mg, Fe, Zn), corrosion and stress shielding phenomena are major areas of concern.83,84,85 Bioceramics, in general, have better compatibility due to their corrosion resistance but their high yield modulus, brittleness and poor workability hinder their efficacy.86,87 In the case of polymeric biomaterials, biocompatibility and processability could be considered their strength and poor degradability, difficulties in sterilization and structural instability as their weakness.80,88

Composite biomaterials have been shown to minimize the limitations of other types of biomaterials along with enhancing their benign aspects.35,38,48,57,72,74,89,90,91,92,93] For hard tissue applications, metallic biomaterials are preferable candidates due to their desirable mechanical properties, which on one hand can compensate for the poor mechanical properties of the additives (ceramics or polymers) and, on the other, provide ease of production for bioactive/biodegradable composites with properties similar to bones.94 For example, even though HAp is biodegradable, it does not have the necessary compressive strength to be used for load-bearing applications. In this case, the addition of Ti to HAp would not only provide the necessary strength but would also provide bioactivity to the prepared implant.83 Similarly, even though magnesium is a desirable candidate for biodegradable implants, the fast corrosion of Mg causes the implant to fail before the bone heals.95 In this case, addition of polymer or ceramic reinforcements can be helpful in controlling the degradation rate and achieving biocompatible corrosion products.96,97

In the following sections, different synthesis routes of metal-based composite biomaterials for hard tissue applications have been illustrated.

Fabrication Routes of Metal-Based Composite Biomaterials

Fabrication of composite biomaterials varies according to the matrix and reinforcement. Thus, in this section, the classification for the fabrication route is based on various combinations of metals, polymers and ceramics.

Surface Functionalization

The biggest problems with metallic biomaterials are their bioinertness (for Ti- and stainless-steel-based implants) and corrosive properties (for Mg- and Zn-based implants). In recent years, several techniques have been developed to make the surface bioactive and corrosion resistant by attaching bioactive ceramics and polymers.98,99,100,101,102,103 As shown in Table II, there are various techniques that can be applied for coating, depending on the desired material to be deposited, coating strength required and investment costs.

Table II Classification of surface modification techniques applied for metallic biomaterials
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Dry Coating Techniques

Various deposition techniques are developed to obtain thin films. Some are based on chemical processes such as chemical vapor deposition (CVD) and molecular beam epitaxy (MBE), and others use physical routes: magnetron sputtering (MS) and pulsed laser deposition (PLD). Due to its numerous advantages, the latter is widely used in surface engineering, for instance, in microelectronics and biomedical applications, and allows easy production of functional coatings.104

These coating methods rely on the high energy provided by the system to achieve fast deposition via melting/vaporization and solidification and/or accelerated flow of particles toward the metallic substrate surfaces.105 These techniques are mostly popular for coating of ceramics in industrial applications. However, recently, these techniques have also been used for biomedical applications. The major advantages and disadvantages of these techniques regarding biomedical applications are shown in Table III.

Table III Some advantages and disadvantages of various dry coating techniques
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Thermal spraying is the most widely used technique for coating of bioceramics on metallic biomaterials. Since the late 1980s, thermal spraying has been a popular technique for coating HAp on metallic biomaterials.106,107,108,109,110 This method uses different media such as flame, plasma arc or electric arc to melt the precursor and spray it onto the surface.111,112 The mechanism of coating includes substrate penetration and/or particle deformation via impact energy.112 The advantages of thermal spraying include rapid deposition and cost-effective coating.113 With the growth in technology, several varieties of thermal spraying, such as flame, arc, atmospheric plasma, vacuum plasma, suspension plasma, coupled radiofrequency plasma, high-velocity suspension flame and gas tunnel plasma spraying, etc., have been implemented for biomedical purposes.114,115,116,117,118 Thermal spraying has also been used to coat implants with HAp/SiO2 composites, HAp/Ti or dicalcium/Ti, Al2O3-TiO2, ZrO2, 45S5 bioglass and more.119,120,121,122,123,124 Thermal sprayed coatings can minimize stress shielding, the major problem in Ti femoral stems used for total hip arthoplasty.125

Currently, the most used biocoatings for implantology are HAp, alumina and titanium nitride. They are frequently coated on Ti and its alloys as well as 316L stainless-steel or CoCr steel alloys. Recently, Durairaj et al. have studied the effect of HAp coating via plasma spraying on anticorrosion properties of metallic biomaterial surfaces.126 The study was conducted for both SS 316L and Ti6Al4V. A better anticorrosion property was found on HAp-coated surfaces for both materials. It was found out that the potential required for localized corrosion increases after HAp coating on both surfaces. The corrosion rate for SS316L and Ti6Al4V was found to be 2.703 mm/year and 0.962 mm/year, respectively, which decreased to 0.0113 mm/year and 0.00801 mm/year using a HAp coating, respectively.

In another study, in vitro and in vivo response of HAp coatings on commercially pure (CP) Ti via coupled radio frequency plasma spraying has been investigated.141 Human fetal osteoblast cells (hFOB) were used to perform cell culture tests for 3 and 11 days, taking CP Ti as reference material. As expected, HAp-coated Ti showed better cell attachment and differentiation compared to uncoated Ti (see Fig. 2). Implantation of the material in cortical defect of rat femur was used to determine the in vivo biocompatibility. The results suggest a higher osteoid content on the HAp-coated Ti implant surface, which is the precondition for better new bone formation, where no osteoid was found on the uncoated Ti surface even after 14 days of implantation, indicating the absence of cellular activity.

Fig. 2
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MTT assay of hFOB cells on Ti- and HAp-coated Ti surfaces as a function of cell culture time. Optical density is directly proportional to the concentration of living cells. A significantly higher concentration of live cells has been found on HAp-coated Ti. *Significant difference in performance. Reprinted with permission from Ref. 141.

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Alumina (Al2O3), which is a polymorphous material,142,143,144 can be deposited by chemical and physical vapor deposition (CVD and PVD).145,146,147,148,149 Furthermore, its chemical inertness150 under physiological conditions associated with strong wear resistance, and excellent hardness makes it an actractive material in tribological151,152 and biomaterial applications.153,154 Joint155 and hip replacement prostheses156 are partly made of Al2O3. For substituing large sections of bone, for instance to replace cancerous bone, porous alumina is prefered because it acts as a scaffold for a new bone growing into the pores.

Titanium nitride (TiN) coatings are often encountered in industrial applications requiring tribological performances coming out from its high surface hardness157 and chemical properties.158 Moreover, TiN is well tolerated by the tissues because of its inertness.159 Furthermore, when TiN is used as an interlayer between HAp films and bulk Ti, the mechanical performances of HAp films are increased because of the improvment in bonding.160 Moreover, PLD161 and MS160 allow producing thin TiN coatings with strong fatigue resistance as well as very strong adherence to the substrate. This results in mechanical properties (e.g., hardness, Young’s modulus, stiffness and mechanical wear) similar to those of bone.

Polymer coating on metallic implants has also been achieved via thermal spraying:162,163,164 In an interesting study, Chebbi et al. were able to achieve a poly 3-hydroxybutyrate-co-3-hydroxyvalerate/polymethylmethacrylate (PHBV/PMMA) and PHBV/PMMA-HAp bi-layer coating on Ti6Al4V implants.107 In vitro analysis of the coated samples suggested that the polymers were able to withstand their bioactive functionalities even after coating, opening possibilities for novel polymer coatings on metallic implants via thermal spraying. Unfortunately, not much research has been focussed in this area in recent years. However, in an inspired research, the vaccum plasma spray (VPS) technique has been shown to coat Ti on polyether ether ketone (PEEK) to improve its biocompatibility along with minimizing the stress-shielding phenomenon.165 In this study, a highly microporous spongy Ti coating was achieved on PEEK with an isoelastic structure, thus preventing spalling of the coating (see Fig. 3). The in vivo study found a significant improvement in bone recovery by the Ti coating, as the rate of new bone formation was found to be much higher for Ti-coated PEEK than for uncoated PEEK (see Fig. 4).This study puts forward a novel path for the preparation of composites for spine implant.

Fig. 3
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Optical microscopy of Ti coating on PEEK. A highly microporous network of coating can be seen. Reprinted with permission from Ref. 165.

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Fig. 4
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New bone formation comparison between control (uncoated PEEK) and Ti-coated PEEK. A significant improvement in bone recovery via Ti coating could be seen. Adapted with permission from Ref. 165.

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Even though widely used, the high temperature applied in thermal spraying is a major problem. The high temperature converts crystalline HAp to αTCP, βTCP, tetra TCP and oxy HAp, which are amorphous in nature. As bone has crystalline HAp, this causes a mismatch between bone and implant, and coated HAp is readily dissolved in blood plasma, causing early instability of the implant and ultimate failure after some time.111,112,166,167,168 In addition, particle size destruction, thermal and residual stresses in the coating, porosity and inhomogenity in the coating are also major causes of concern in using this technique.169,170 The mechanical properties of the HAp coating via thermal spraying are also dependent on the thickness of the coating. A study found that a HAp coating of about 150 µm via VPS on Ti6Al4V significantly diminished the fatigue strength of the coating, which was not the case in 25–100-µm-thick HAp coatings.171 Moreover, thermal spray coating implant devices with HAp on metal substrates showed disadvantages and failure at interface due to the compressive residual stresses occurring during the coating.170

To minimize the above-mentioned high-temperature problems, several studies on the application of cold spraying for biomedical applications have also been done.172,173,174,175 In cold spraying, the particles are heated below its melting temperature and accelerated to very high speed using a carrier gas toward the substrate to achieve adhesion.169 It provides an outstanding advantage of room temperature deposition with high deposition efficiency, where the particles can retain their crystal structure even after deposition.169,176 Its ease of application has allowed the successful coating of HAp, HAp/Ti, HAp-Ag/PEEK and HAp/graphene composites on various metallic substrates with intact structural features.131,177,178,179,180

In a recent study, the efficiency of cold spraying to coat Ti6Al4V alloys with tantalium (Ta) and its in vitro bioactivity has been investigated.181 A rough and porous coating of Ta was achieved without oxidation (see Fig. 5), which was able to promote HAp mineralization in SBF solution, resulting in good bioactivity. Literature suggests that rough surface is beneficial for the cell attachment and proliferation.182,183 Thus, this study suggests a good path for bioactivation of inert metallic implants.

Fig. 5
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(a) Macroscopic, (b) laser confocal and (c) and (d) SEM images of the surface morphology of tatalium (Ta) coating via cold spraying. A rough and microporous surface can be seen. Reprinted with permission from Ref. 181.

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Other techniques have also been applied for biomedical applications to minimize or mitigate these problems. PVD-based deposition techniques are one of the best alternatives to thermal spraying. Sputtering, PLD, ion beam-assisted deposition, arc ion plating and cathode arc deposition are the most popular PVD-based techniques for biomedical applications. Among these, sputtering is the most commonly used PVD-based technique where coating is achieved by ejecting material from the “target,” i.e., the material to be deposited, via ion bombardment. These ejected particles later move toward the substrate and form a coating via condensation in a reactive or inert environment.184 This technique allows a better controlled deposition of thin coatings.185 The adhesion strength of coating achieved by sputtering was found to be higher than that of thermal (plasma) spraying, given its cleanliness, enhanced atomic diffusion and mixing at interface as well as better mechanical interlocking.113,186 Given these advantages, it has been used for coating metallic substrates with HAp, CaP, Si-HAp, Sr-HAp, F-HAp, Mg-HAp, Ag-HAp, Ag-SiC-HAp, Ti-Si-N, Ti-N along with others.128,187,188,189,190,191,192,193,194,195 In a recent study, Kumar et al. applied a coating of transitional metal carbides, i.e., ZrC and TiC, via sputtering to improve the bio-corrosion and antibacterial properties of stainless-steel (SS) 316L.196 The coated surfaces provided better anticorrosion properties in an artificial blood plasma (ABP) solution. The coating also proved to have better antibacterial properties. The bacterial adhesion test results clearly provide an overview of the bacterial adhesion and biofilm formation on SS 316L samples, which was absent on TiC- and ZrC-coated SS 316L samples (see Fig. 6).

Fig. 6
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Epifluorescence microscope images after 4 days of incubation against P. aeruginosa of (a) uncoated SS 316L, (b) TiC-coated and (c) ZrC-coated SS 316L substrates. The bacterial adhesion and biofilm formation can be seen for uncoated SS 316L substrates, which was prevented by the coating of TiC and ZrC. Reprinted with permission from Ref. 196.

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The amorphous nature of the coating is also a problem in the case of sputtering. However, with the help of heat treatment, a crystalline coating could be achieved.197,198,199 Hamdi et al. illustrated the improvement in adhesion of sputtering-deposited coating via heat treatment.198 In this study, a thin Ag-TaO film was deposited on SS 316L via magnetron sputtering, which was initially found to be amorphous in nature. With the help of heat treatment, a tremendous improvement in crystallinity and smooth coarsening was obtained (see Fig. 7). A significant enhancement in mechanical properties was also found after heat treatment. Where the scratch distance and critical load was found to be 440 µm and 672 mN for the pretreatment coating, it improved to 757 µm and 2749 mN, respectively, after annealing at 500°C for 60 min.

Fig. 7
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FESEM surface morphology and cross-section view of Ag-TaO film. a and c Surface morphology of the sputtered film before and after heat treatment. b and d Cross-section of the sputtered film before and after heat treatment. Reprinted with permission from Ref. 198.

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Another popular PVD-based deposition technique for biomedical application is pulsed laser deposition (PLD). In this technique, high-powered laser energy is used to vaporize the to-be-coated material and subsequently deposit on the substrate.200 One of the distinct advantages of this technique is its ability to preserve the stoichiometry of the coating after deposition.201,202 Given this advantage, a very dense and crystalline coating of HAp is possible via PLD.203,204,205 This is evident in the comparitive study of García-Sanz et al., performed between HAp coating on Ti produced by plasma spraying and PLD.206 Where plasma spraying produced a thick, brittle and inhomogeneous HAp coating, PLD proved to be a better alternative by producing a well-adhered, non-brittle and homogeneous thin coating. However, a low deposition rate, splashing and higher costs have always been major drawbacks to PLD.207 Therefore, several variations of PLD such as pulsed electron deposition (PED) and pulsed plasma deposition (PPD) are also being currently investigated for biomedical applications.129,208,209,210,211

Bianchi et al. investigated the nano-mechanical and in vitro properties of CaP coatings on CP Ti grade 2 processed via PPD technology.208 To achieve a crystallized CaP coating, annealing of the samples was performed for 1 h at 600°C. The micro-scratch test results clearly showed the difference in coating adhesion of non-annealed and annealed samples; where spalling/wedging was observed in non-annealed samples, this was absent in the case of the annealed CaP coating and ductile tensile cracking was observed. Interestingly, in the case of osteoblast cell adhesion and proliferation, no significant difference was observed in annealed and non-annealed CaP coatings after a 3-day cell culture. This study suggests that the amorphous coating problem in sputtering also occurs in PLD technology.

Another interesting high-temperature coating technique for biomedical applications is hot isostatic pressing (HIP). This process applies high isostatic pressure at high temperatures in an encapsulated system by means of gas to achieve a fully densified ceramic coating.113,130 The application of uniform pressure all over the system makes it better than other uniaxial pressing techniques such as hot pressing, as this helps in removing the shape limitation of the substrate for coating. Using HIP, coating of HAp at a relatively low temperature of 135°C with a relative density close to that of bone is possible.212 In another study, HIP also proved to be a viable technique for coating partially stabilized zirconia (PSZ)/HAp composites.130 The low temperature applied in HIP along with the encapsulated system prevents thermal degradation of HAp, providing it a unique advantage compared to other high-temperature coating techniques.130

Wet Coating Techniques

Covalent Bonding

Recently, covalent grafting of polymer chains to metal surfaces has been proposed to improve the adhesion, long-term stability and durability between the metal surface and the polymer coating.223,224,225

Thus, many efforts have been made to synthesize grafted polymers to control the polymer thickness and the adhesion strength.226,227 Two principal methodologies have been used to link polymer chains on metal surfaces: the “grafting to” or “grafting from” techniques.228,229,230

“Grafting to” requires the synthesis of end-functionalized polymers, i.e., polymer chains terminated with functional groups and their adsorption, through either physical or covalent bonds, to functionalize surfaces.231 However, although it makes it possible to characterize the polymer before the grafting reaction, its effectiveness diminishes with the heaviness of the polymer. In fact, the steric repulsions generated by the anchored chains may impede the diffusion of the macromolecules toward the surface, preventing the arrangement of thick and opaque polymer layers.231,232

The “grafting to” method allows grafting known structures and often needs an anchor molecule (see Fig. 8). The synthesis of the well-defined polymer was achieved by primarily applying living polymerization techniques, including controlled radical polymerization (CRP).233,234,235 Of these living radical polymerization techniques, three are the most well known: nitroxide-mediated polymerization, atom transfer radical polymerization (ATRP) and reversible addition-fragmentation chain transfer (RAFT) polymerization. These procedures for covalently tying well-defined polymer onto surfaces have been developed, counting the covalent connection of end-functionalized polymers incorporating an appropriate anchor (“grafting to”) and the in situ polymerization initiated from the surface (“grafting from”).235 For example, the most common techniques for the covalent connection of target molecules onto titanium surfaces include the formation of the monolayer with organofunctional silanes,236,237 phosphonic acids238,239,240,241 and phosphonates.242 Then, a target molecule was covalently joined to the surface. However, these techniques usually require different steps to create suitable organofunctional anchors. Recently, the catechol unit has become a fundamental biomimetic ligand for surface immobilization.243 Inspired by the interesting properties of the adhesive proteins secreted by marine life forms, consolidating the catechol functionality into macromolecules has raised much interest.244,245,246 Messersmith et al.245 have detailed the first example of anchoring a catechol initiator onto stainless-steel surfaces for consequent surface initiated ATRP. In this study, the catechol units have been conjugated with different polymerization initiators or “clickable” moieties.134,247 For illustration, Woisel et al.245 report a flexible and precise approach for titanium surface modification based on catechol surface functionalization that enables target molecules to be attached utilizing the copper-catalyzed azide-alkyne cycloaddition reaction. Chouirfa et al. have developed a simple and flexible methodology based on catechol surface modification that was created to empower bioactive ionic polymers bearing thiol end groups to be attached employing a thiolene click reaction.247,248

Fig. 8
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“Grafting to” process.

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The “grafting from” method involves the polymer chains’ growth from initiator-modified surfaces through surface-initiated polymerizations (see Fig. 9).249

Fig. 9
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“Grafting from” process.

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This enables less demanding diffusion of monomer molecules toward the propagating radicals fixed on the surface. This infers a higher graft density, allowing denser, thicker and more stable organic layers with better control of the polymer thickness.249,250

Non-covalent Bonding

Electrophoretic deposition Electrophoretic deposition (EPD) is a simple and versatile colloidal solution-based coating technique for coating at room temperature. It consists of two steps: first, charged particles are moved under the influence of an electric field to the opposite charged electrode. Following this step, the charged particles are deposited via coagulation and precipitation on the electrode surface to form a coating.251 EPD is quite a beneficial coating technique for biomedical applications, given to its simple equipment requirements, short processing time, good control in coating thickness, ability to produce interconnected porous coatings and ease of coating even on complex-shaped structures.251,252,253,254,255 Given to these advantages, not only biocompatible ceramics256,257,258 and polymers,259 but also a wide variety of composites have been deposited in metallic implants.260,261,262,263,264

In a recent study, PMMA and PMMA-alumina composites have been coated on stainless steel via EPD to improve the bioactivity and osteoconductive properties.265 They were able to achieve a homogeneous and porous coating in both cases, where PMMA coating had maximum thickness of 2 µm and PMMA-alumina had a coating thickness of 10 µm. The ALP activity studies confirmed the bioactivity and osteoconductive capacity of both coated surfaces, where maximum bioactivity was found for composite-coated surfaces followed by PMMA-coated surfaces.

Physisorption

From a biological perspective, a hydrophilic surface provides a favorable environment for cell growth and osteogenesis by improving protein interaction with the surface.266,267,268 This hydrophilicity can be achieved via physisorption. It is a relatively simple and versatile technique that does not require complex equipment. In this technique, metallic specimens can be directly immersed in the desired polymer solution to achieve polymer physiosorbed metallic substrates. Migonney et al. have used this technique to illustrate how sulfonate and/or carboxylate groups bearing polymers such as PolyNaSS and PAA can be easily physiosorbed on Ti6Al4V surfaces.137

Sol–Gel

To deposit HAp on metal substrate devoted to implants, plasma spraying,269,270 dipping,271 electro-co-deposition,272 PLD,273 sputtering274 and sol-gel-derived coating275,276,277 have been generally used. Nevertheless, due to the high-temperature process,278 plasma spray leads to a lack of crystallinity and the modification of HAp composition.

However, sol-gel is a very versatile and easy method, involving simple fluid flow and evaporation process and generating rather uniform coatings].279 Moreover, sol-gel HAp coatings can be produced at temperatures as low as 900°C, thus avoiding the decomposition of HAp and Ti alloy phase transformation.

Mechanical and Wear Properties of Coatings

Bioceramics are able to directly bond with the surrounding tissue once implanted.280 However, typically the principal weakness of these materials originates from the low mechanical strengths making them inappropriate for load-bearing applications.

The mechanical properties and tribological behaviors of Al2O3 and HAp have been studied owing to their microstructural features.281,282,283,284 On the one side the wear mechanisms161 and on the other the micro-hardness (H) and elastic modulus (E)285 were described using nano-scratch and nano-indentation tests, respectively. The values of H and E experimentally determined in Ref. 285 are in good agreement with the data from other studies.286,287 Most of the reported H and E values for HAp coatings on Ti are between 83 GPa and 123 GPa and between 4 GPa and 5 GPa, respectively.288

The combination of metallic implants and bioceramic thin films brings the advantages of the metal with its mechanical properties and those of the ceramic layer with its bioactive behavior. The mechanical properties are preserved while the integration of the prosthesis with the freshly grown bone is improved. Nonetheless, the metal-ceramics interfaces are frequently the seat of residual stresses.289 Their magnitude predominantly depends on the deposition processes and the coating conditions, which can result in high stress and cause a delamination at the interface.161,290,291,292 In Ref. 140, the first mechanical approach to residual stress was reported in a Ti substrate before and after coating. To obtain accurate residual strains measurements, x-ray diffraction by sin2ψ method is the well-adapted technique to overcome problem links to high surface roughness and < 1 mm layer thickness.

Liquid State Processing

Due to the economical nature and simplicity of these techniques, they are widely used for metal matrix composite (MMC) preparation. They can be mainly divided into two parts: casting and infiltration.

Casting

This is one of the most traditional techniques used for manufacturing complex-shaped MMCs in an economical manner. In this process, at first reinforcements are added in liquid metal while continuously stirring to achieve a homogeneous distribution. Liquid material is afterwards poured into a mold with the desired cavity shape and then allowed to solidify. The solidified part is later ejected or broken out of the mold, providing the final part.293

Given to its simplicity and efficiency, stir casting is the most common casting technique applied for preparation of metal-matrix composites. In this process, the reinforcements are incorporated into the melt and distributed into the melt via mechanical stirring. The casted composites can be further extruded to reduce the porosity and homogenize the reinforcement distribution.294 This technique has been used to incorporate various biomaterials on magnesium and its alloys.70,295,296,297,298 Khanra et al. have used stir casting and extrusion process to reinforce Mg and Mg-6Zn-1Mn (ZM61) with 5, 10 and 15 wt.% HAp.295 The addition of HAp reinforcements led to an increase in hardness at the expense of ductility. The non-uniform distribution of HAp in the matrix led to lower experimental hardness than theoretically estimated values and irregularities in the tensile strength of composites.

A further improvement in stir casting technique for obtaining a uniform distribution of reinforcement in composite is the high shear solidification technique. In this technique the liquid metal is subjected to intensive shear via high shear rotor-mixer leading to better mixing of reinforcement in metal matrix.299,300,301,302 Huang et al. have employed this technique with the equal channel angular extrusion (ECAE) process to successfully develop Mg-HAp nanocomposites.302 They initially melted Mg and added Zn and Zr to produce Mg-3Zn-0.5 Zr alloy in which 1 wt.%, 3 wt.%, 5 wt.% and 10 wt.%. HAp was added. Afterwards, the ECAE process was performed at a temperature of 300°C for several cycles. With the increase in HAp wt.%, a decrease in grain size was found, where the HAp particles were found to be coupled to the alloy grain-boundary network. In SEM analysis, the individual aggregate of HAp particles was found to be in submicron scale, where some aggregates consist of even only a few HAp particles (see Fig. 10). This shows that the HAp particles were loose and as result were filled and penetrated by matrix alloy during shear mixing. The ECAE process resulted in further grain refinement.

Fig. 10
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SEM images of electropolished Mg-3Zn-0.5Zr-5HAp composite at different magnifications. (a) Loose HAp particle aggregates, (b) typical dimensions of these aggregates and (c) fine HAp aggregates containing few HAp particles. Reprinted with permission from Ref. 302.

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Infiltration

In this technique, molten matrix metal is impregnated in porous preforms of desired reinforcements, where molten metal fills the pores with or without the application of external force.303 It has the advantage of preparing MMCs with uniform distribution of reinforcements, eliminating residual porosities and near-net-shaped composite production.304

Several interesting studies have been performed to prepare Mg matrix composites using this technique.305,306,307,308,309 Ma et al. have applied this technique to prepare a biodegradable β-TCP/Mg-Ca composite.306 In this study, they initially created the interconnected porous β-TCP scaffolds by coating PU foams with β-TCP via slurry coating and sintering at 1100°C to burn out PU foam. The prepared scaffolds were infiltrated by Mg-1 wt.% Ca alloy at a temperature of 700–720°C via suction casting. Excellent wettability between β-TCP and Mg alloy resulted in full penetration of alloy throughout the porous scaffold (see Fig. 11). Addition of Mg also led to a tremendous improvement in the compressive strength of the composite, where the ultimate compressive strength of the final composite (147 ± 3 MPa) was found to be close to that of a bone (180 MPa). The anticorrosion properties of the composite were also found to be better than that of the Mg-Ca alloys.

Fig. 11
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SEM images of (a) PU foam, (b) porous β-TCP scaffold, (c) β-TCP/Mg-Ca composite and (d) the interface between the β-TCP scaffold and Mg-Ca alloy. The Mg-C-β-TCP interface had no discernable debonding or microcracks, where the Mg-Ca alloy was also able to penetrate the hollow section of the scaffold. Reprinted with permission from Ref. 306.

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In a recent study, Myalska et al. have infiltrated a carbon foam with pure Mg, Mg-3Al-1Zn (AZ31) and Mg-Zn-RE-Zr (RZ5) Mg alloys in an attempt to develop a new composite material.308 They have used the open-celled rectangular carbon foam (Cof) and pressure infiltration technique. In their study, composite prepared using Mg and AZ31 alloy showed a good infiltration tendency; this was not found to be the case for RZ5 alloy-based composite. The rare earth elements present in RZ5 alloy had a destructive influence on Cof, as along with micro- and macro-pores, carbon foam damage was also found in this case. Best infiltration was found for Cof-Mg composite with a porosity of only 1.4%, which was slightly worse in the case of AZ31 alloy with a porosity of 2.4%. The reinforcement of carbon foam led to an improvement in compressive strength and stiffness for Mg and the AZ31 alloy.

Solid-State Processing

Solid-state processing mostly uses diffusion bonding or sintering for the preparation of metal matrix composites. Processes that use one of those principles for their purpose follow.

Powder Metallurgy

Powder metallurgy (PM) is the most common route for the development of MMCs. It has the advantage of producing composites with uniform distribution of reinforcements in matrix, requires less temperature and is cost-effective.310 It mainly consists of three steps: powder blending, die compaction and sintering. Due to its simplicity and effectiveness, several metal matrix composites have been prepared using this technique for biomedical applications.311,312,313,314,315,316,317

The final porous structure of the product via this technique is especially helpful in the case of Ti-HAp composites, as it can help in minimizing the elastic modulus and provide osseointegration.314 To evaluate this, Ning et al. have performed in vivo and in vitro tests on HAp-Ti composites.314 They have prepared 30 wt.%, 50 wt.% and 70 wt.%. Ti-HAp composites via powder metallurgy, sintering at 1200°C, using 10 wt.% bioactive glass as a sintering aid. They have used SBF to evaluate the apatite formation characteristics of the composites. They have found a good correlation between the Ti content in composite to the bioactivity, as apatite formation occurred for the case of 50 and 70Ti-HAp samples, which was not the case for 30Ti-HAp samples. It was suggested that the formation of α Ti and Ti2O on 50 and 70Ti-HAp samples was responsible for apatite formation. For the in vivo study, they implanted the composites on the metaphases of rabbit femur. Where after a month of implantation, bone growth was found to be slow on 30Ti-HAp implants compared to their counterparts, after 6 months of implantation all the Ti-HAp composites formed a good bone-bonding interface with host bone via apatite layer.

Even though HAp is stable up to 1250°C in controlled environment, in Ti-HAp systems the presence of Ti leads to accelerated dihydroxylation and decomposition of HAp at temperatures > 800°C forming tetracalcium phosphate and calcium oxide.318,319 Thus, to minimize this effect, Comin et al. have prepared Ti-20HAp composites using PM, sintering at a temperature of 800°C in argon environment for 2 h.312 XRD analysis showed that HAp and Ti existed in their simple forms, suggesting no reaction between HAp and Ti took place during the manufacturing process. The composites have shown great bioactivity and cytocompatibility. After 10 days in SBF solution, the whole surface was found to be covered with apatite. In cell culture study, NIH3T3 cells were able to adhere and proliferate to the whole surface, even to the pores by day 4 of cell culture.

Along with Ti, enhancement in Mg-based implants has also been attained via this technique. Khalajabadi has applied this technique to provide the beneficial properties of HAp and TiO2 to enhance the biodegradability of Mg-based implants.315 The prepared Mg/x1HAp/x2TiO2 (x1 = 27.5, 20, 12.5, 5 and x2 = 0, 5, 10, 15 wt.%, respectively) composite via uniaxial pressing of powder mixture at ~ 840 MPa in RT followed by sintering at 400°C for 1–2 h in argon environment. The ultimate compressive strength (UCS) and compressive failure strain (CFS) were investigated for all the composites. The highest UCS of ~ 237 MPa was found for Mg/27.5HAp samples, where the addition of 5 wt.% TiO2 and decreasing the HAp to 20.% led to a decrease of UCS to ~ 188 MPa but the CFS had improved from 1.7% to 1.9%. Interestingly, decreasing the HAp to 5 wt.% and increasing the TiO2 to 15 wt.% led to dramatic improvement in UCS from 1.9% to 8.1%. With higher content of HAp, the agglomeration of HAp particles in the composite causes difficulty in deformation by hindering dislocation and hence increase in UCS. The homogeneous distribution of TiO2 led to a decrease in stress concentration and voids near big particles, which helped in improvement of CFS. The addition of TiO2 particles also helped in improving the corrosion behavior by decreasing the number of pores and voids around HAp agglomerates.

In the latest addition, additive manufacturing (AM) technologies utilizing powders as the feed material are taking powder metallurgy to the next level by applying its principle to produce custom-shaped metal-matrix composites. Selective laser melting (SLM) is the most commonly used AM technology used for producing metal-matrix composite for biomedical applications.67,320,321,322 In this technique, the initial 3D model is provided to machine using CAD software. Afterwards, the powder is laid on the molding platform, which is selectively melted using high-energy laser beams in an inert environment and subsequently solidified, where a new layer is added in a step-by-step fashion creating the final structure.323 For production of metal-matrix composites, reinforcements can be incorporated ex situ, where the reinforcement is mixed in metallic powder and subsequently the powder mixture is used for additive manufacturing.324 The other way to produce a metal-matrix composite is in situ where the reinforcement is produced in the matrix by chemical reaction between metal matrix and elements or compounds.325

To provide better wear and hardness properties to Ti-based composites, TiB has been added as a reinforcement in the production of Ti-TiB composites via in situ SLM.326 TiB particles were introduced to the Ti matrix by an in situ reaction between Ti and TiB2 during the SLM process, which provided good interfacial bonding between matrix and reinforcements. The TEM analysis identified a significant grain refinement of the α-Ti grains due to the presence of B and the rapid solidification. A needle-shaped structure of TiB was seen in SEM images due to the directional growth of the TiB by the influence of B-B bonds, which are stronger in one direction (see Fig. 12). The addition of TiB led to an improvement in tensile, hardness and compressive strength of the composite.

Fig. 12
figure 12

SEM images of the SLM-produced Ti-TiB composite microstructure at different magnifications: (a, b) cross-sectional views; (c, d) longitudinal views. White arrows indicate TiB particles where the needle-shaped TiB particles within the Ti matrix can be seen. Reprinted with permission from Ref. 326.

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In a recent study, a biodegradable Mg alloy/BG composite has been prepared by SLM process.321 In this study, ZK30 (Mg-3Zn-0.6Zr) and BG powders were mixed via ball milling and fabricated via SLM to prepare ZK30/xBG (x = 0, 5, 10, 15 wt.%) composites. With the addition of higher content of BG, the number and size of pores increased in the composite. In microhardness tests, the hardness of composite increased with the addition of BG but the increase was found to be inhomogeneous. In the degradation rate test, the addition of BG led to an improvement in hydrogen evolution and corrosion potential of the composites (see Fig. 13). However, the excessive content of BG led to more defects, which accelerated the degradation rate. In cytotoxicity analysis, the increase in BG content enhanced the cytocompatibility as the relative growth rate of L929 cells was higher in ZK30/10BG and ZK30/15BG compared to ZK30 and ZK30/5BG composites (see Fig. 14). It was assumed that the reduction in release of Mg2+ and increase in bioactivity by higher content of BG could be responsible for this behavior.

Fig. 13
figure 13

Hydrogen evolution (a) and Tafel curves (b) of ZK30/xBG composites. Increasing the BG content, the hydrogen evolution and corrosion potential decrease if the BG content does not become too high, i.e., ZK30/15BG. Reprinted with permission from Ref. 321.

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Fig. 14
figure 14

Relative growth rate of L929 cells on ZK30/xBG composites after (a) 24 h, (b) 48 h and (c) 72 h. A higher rate in ZK30/10BG and ZK30/15BG can be seen compared to other samples on all days. Adapted with permission from Ref. 321.

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Pressure-Assisted Sandwich Formation

The interdiffusion between polymers is applied for the preparation of metal/polymer composite in this technique. Metal/polymer composites are generally prepared via accumulative roll bonding in the automotive and aerospace industry where resins are used as adhesion agent327,328,329,330. However, resins are found to be cytotoxic in nature, which limits this process for biomedical applications.331 To overcome this technique, an innovative approach is developed by Reggente et al. where grafted polymers on metal sheets were used as an adhesive agent to prepare Ti/PMMA/Ti sandwich materials with tunable mechanical properties.332 In this study, PMMA chains were initially grafted on Ti sheets using the “grafting from” technique.333 Afterwards, PMMA-grafted Ti sheets and PMMA sheets were pressed together via hot-pressing to prepare Ti/PMMA/Ti sandwiches. The bonding strength was investigated via pull-off and shear tests. The best bonding conditions, i.e., pull-off strength of ~ 20 MPa and ultimate shear strength of ~ 10 MPa, were obtained at hot-pressing at a temperature of 180°C and a pressure of 0.2 MPa for 90 min. The mechanical properties of these sandwiches are easily tunable by varying the PMMA/Ti sheet thickness or quality of the metal. The experimentally obtained tensile properties seemed to correlate with the theoretical values obtained using the rules of mixture (RoM) (Table IV). The Young’s modulus of these sandwiches (~ 51 GPa) was quite close to that of cortical bone; thus, they could minimize the stress shielding problem commonly seen for Ti implants. The mechanical properties of these sandwiches can be easily reached closer to bone by increasing the ratio of the polymer if keeping the metal sheets’ thickness constant. The result obtained from 3-point bending, deep drawing and Erichsen tests of these sandwiches suggested an excellent possibility to shape these sandwiches to fulfill patients’ needs.

Table IV Tensile properties of Ti, PMMA and Ti/PMMA/Ti sandwiches and their theoretically estimated values via RoM.
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Thus, this technique can be a new alternative to obtained custom-shaped composites where mechanical properties can also be tuned based on its application area.

Conclusion and Outlook

The purpose of the current review is to provide a short summary of the possibilities to obtain metal-based composites that can be of use for hard-tissue applications. The most used and emerging techniques have been mentioned in this review. Based on necessity, funds and application area, both biodegradable and bioactive composites can be manufactured. When the need is just to obtain a strong bonding between the surface of implants and the surrounding tissues, coatings seemed to be the way to go. Although thermal spraying is the most used technique in the past, PVD is emerging quickly, given its ease of control and high-quality coatings. Coatings can also be provided via wet techniques in an energy-friendly way.

If the necessity is to develop metal-matrix composites, both liquid-state and solid-state processing could be applied. However, solid-state processing techniques such as PM seemed to provide better end results by providing fine-grain-sized porous microstructures enhancing the cell interaction with the implants. AM is currently the most researched technique for biomedical applications as it minimizes or in some case stops the need for shaping after the manufacturing process by providing custom-shaped parts. Pressure-assisted sandwich formation could also be a promising processing technique in the future as it provides ease of production.