**5. Surface modification techniques**

#### **5.1. Sol-gel**

The sol-gel method has been widely used to deposit calcium phosphate onto dense or po‐ rous metallic materials. There are two routes for a sol-gel reaction, namely inorganic and or‐ ganic, using reagents consisting of a colloidal suspension solution of inorganic or organic precursors. The sol-gel technique transforms a liquid (sol) into a solid phase (gel) and re‐ quires drying and heat treatment stages. The advantages of the sol-gel method include: (i) it is cost-effective, (ii) it is easy to control the final chemical composition and thickness of the coating, (iii) the coating is readily anchored on the substrate, and it is usually homogenous with a good surface finish, and (iv) it can be used for coating implants or substrates that have complex surfaces or large surface areas.

Wen *et al.* [22] reported that a sol-gel method for HA and titania (TiO2) coatings exhibited excellent bioactivity after immersion in simulated body fluid (SBF) that mimics human body fluid of a similar ion concentration and pH value to human blood plasma. In addition to en‐ hancing titanium bioactivity, HA-titania coating is expected to increase the bonding strength and corrosion resistance. The surface morphology and microstructure of HA and titania coating before and after being immersed in SBF are presented in Figure 1 (a)-(d). It can be seen that the coating is dense, uniform and without cracks. Wen *et al.* also reported that after soaking in SBF, HA granules grow gradually.

**Figure 1.** SEM micrographs of the surface morphology of HA/TiO2 films after soaking in SBF for (a) 0 day, (b) 1 day, (c) 8 days, and (d) 15 days

#### **5.2. Electrodeposition of materials**

bone apatite and collagen production [15-16]. It is suggested that altering the nanostruc‐ tured surface morphology influences the apatite inducing ability and improves osteoblast

Calcium phosphate is a synthetic ceramic that has been proven to support bone apposition and to enhance the osteoconduction of the bone. Calcium phosphate ceramics for bone tis‐ sue applications include tricalcium phosphate (TCP), octocalcium phosphate (OCP), hydrox‐ yapatite (Ca10(PO4)6(OH)2, HA), and biphasic calcium phosphate (BCP) [18]. These ceramics accelerate the healing process and have been widely used in conjunction with metallic mate‐ rial as a bioactive coating material. The ratio of Ca/P in calcium phosphate should resemble the biological apatite mineral of bone (*i*.*e*., 1.50-1.69). Calcium phosphate has the natural fa‐

Hydroxyapatite demonstrates the best bioactivity amongst all the forms of calcium phos‐ phate. Hydroxyapatite (HA) exhibits functionality in promoting osteoblast adhesion, migra‐ tion, differentiation and proliferation; all of which are essential for bone regeneration. HA also has the ability to bond directly onto bone. The bioactivity of HA has made this ceramic the favourite for implant applications. HA nanoparticles may also induce cancer cell apopto‐ sis [19]. The crystalline form of HA exhibits biointegration and prevents formation of ad‐ verse fibrous tissue. It is a more desirable coating than amorphous HA due to its ability to provide a better substrate for a different cell line [20]. Amorphous HA tends to dissolve in human fluid more easily and leads to loosening of the implant. Nanocrystalline HA is more favourable than microcrystalline HA because of its structural similarity with apatite [21].

The sol-gel method has been widely used to deposit calcium phosphate onto dense or po‐ rous metallic materials. There are two routes for a sol-gel reaction, namely inorganic and or‐ ganic, using reagents consisting of a colloidal suspension solution of inorganic or organic precursors. The sol-gel technique transforms a liquid (sol) into a solid phase (gel) and re‐ quires drying and heat treatment stages. The advantages of the sol-gel method include: (i) it is cost-effective, (ii) it is easy to control the final chemical composition and thickness of the coating, (iii) the coating is readily anchored on the substrate, and it is usually homogenous with a good surface finish, and (iv) it can be used for coating implants or substrates that

Wen *et al.* [22] reported that a sol-gel method for HA and titania (TiO2) coatings exhibited excellent bioactivity after immersion in simulated body fluid (SBF) that mimics human body

adhesion and differentiation [17].

26 Titanium Alloys - Advances in Properties Control

**4.1. Calcium phosphate coatings**

cility to bond directly to bone.

**4.2. Nano-hydroxyapatite coatings**

**5. Surface modification techniques**

have complex surfaces or large surface areas.

**5.1. Sol-gel**

Electrodeposition is a coating method applied to the fabrication of computer chips and mag‐ netic data storage. Recently, that has been rising interest in electrochemical deposition for tissue engineering applications due to its ability to coat complex 3D components.

Lopez-Heredia *et al.* [23] coated calcium phosphate onto porous titanium using the electro‐ deposition method. In the process, Ti, platinum mesh, and supersaturated calcium phos‐ phate solution were used as the cathode, electrode and electrolyte, respectively. The ratio of Ca/P in the calcium phosphate coating was 1.65 and the coating thickness was 25 µm. The calcium phosphate coating was homogenous and covered the entire Ti surface. Moreover, they reported that the coating presented good adhesion to the underlying substrate. The electrodeposition of CaP showed that calcium phosphate enhances the adherence of cells.

**5.4. Thermal spray**

bone or dental applications.

stock particle size and velocity.

the same as the starting materials.

**5.5. Physical vapor deposition**

must be controlled to produce a high quality coating.

The thermal spray technique is a well-established and versatile technique that can be ap‐ plied for a wide variety of coating materials, *i*.*e*., metallic, non-metallic, ceramic, and poly‐ meric. Thermal spray coated medical implants, such as orthopaedic and dental prostheses, have been commercially used. Thermal spray offers several advantages, such as the ability to coat low and high melting materials, a high deposition rate, and flexibility in coating 3D shape components. It is also cost effective [28]. Despite these advantages, some problems have been revealed after long term implantations using thermal spray coatings, such as de‐ lamination, resorption, biodegradation of the thick coating and mechanical instability [29]. Thus, improving the adhesion strength of thermal sprayed coatings is a major concern for

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There are several types of thermal spraying; for example, plasma spray, flame spray, and cold spray [29]. Plasma spray is commonly applied to produce thick coatings for metallic corrosion protection. It is also flexible, due to its ability to coat different substrates. During plasma spraying, the precursor is atomised and injected into plasma jet, then accelerated to‐ wards the substrate with the aid of an inert carrier gas [30]. There are many parameters that

Flame spray uses a combustion flame to melt the solid precursor. There is, additionally, an‐ other type of flame spray termed as high velocity oxygen fuel (HVOF). This technology is favourable due to its high spray velocity and the formation of a strong bond coating [30-32]. The thermal spray technique has been widely employed for HA coatings. The surface morphology of HA coatings obtained with various parameters of stand-off distance (SOD) and power are presented in Figure 2 (a)-(d). Sun *et al.* [28] reported that when the spray power increased, the crystallinity of HA decreased and the amorphous phase became more obvious. The effect of SOD indicated an inverse correlation with deposition efficien‐ cy. Several parameters that influence the deposition of HA are SOD, spray power, feed‐

Cold spray is a new member of the thermal spray family. This technique uses small particles of 1-50 µm. A supersonic jet of compressed gas is used to accelerate the particles. The ad‐ vantage of using this technique is the ability to produce dense coatings and maintain the material chemistry and phase composition of the feedstock. Noppakun *et al.* [33] have ap‐ plied cold spray technique to deposit HA-Ag/poly-esther-ether-ketone on glass slides. This study reported that cold spray was able to retain and elicit a coating functionality that was

Physical vapor deposition (PVD) is a deposition method where materials are evaporated or sputtered, transferred and deposited onto the substrate surface. This physical process in‐ cludes thermal evaporation or plasma-induced ion bombardment onto the sputtering target. A condensation or reaction of the coating materials then takes place on the substrate surface to form coatings. Variants of the PVD process include evaporation, ion plating, pulsed laser dep‐

Adamek *et al.* [24] succeeded in producing an HA coating on porous Ti6Al4V using electro‐ deposition. The intermediate layer between the bone and metallic implant was rough and porous. Large pores and nanolamellae were present within the HA layer. The flexibility of the electrodeposition technique for coating solid and porous metallic implants has acquired increasing interest due to this ability to enhance the bioactivity of bone implant materials.

#### **5.3. Biomimetic creation of surfaces**

There are two major steps involved in the biomimetic technique. The first step is to conduct a pre-treatment of the implant material surface to create a layer functional group that can induce formation of an effective apatite layer. Several studies have revealed that an apatite layer has not been formed on materials without any treatment prior to immersion [25]. Pre‐ liminary treatment includes, for example, hydrothermal, sol-gel, alkali heat treatment and micro-arc. The second step is to immerse the biomaterials into a simulated body fluid (SBF). In this step, the bone apatite layer is formed on the biomaterial's surface. The high apatite forming ability of titanium arises from the formation of a hydrated titanate surface layer during chemical treatment. The advantages of the biomimetic process include (i) flexibility in the control of the chemical composition and thickness of the coating, (ii) the formation of relatively homogenous bioactive bonelike apatite coatings, (iii) a lower processing tempera‐ ture, and (iv) the ability to coat 3D geometries.

Wang *et al.* [25] used a modified biomimetic approach to improve the biocompatibility of porous titanium alloy scaffolds. In their experiment, porous Ti10Nb10Zr underwent an alka‐ li heat treatment prior to soaking in SBF. Two NaOH concentrations of 5 M and 0.5 M were used, and the samples were soaked for 1 week. The surface morphologies of porous TiNbZr after alkali soaking and heat treatment revealed a nanofiber layer, that consisted of sodium titanate. Parameters that influenced the morphology and thickness of the sodium titanate were reaction temperature and NaOH concentration.

Calcium phosphate was successfully deposited on the surface of the porous TiNbZr. The calcium phosphate layer was uniform and homogenously spread onto the surface. Anoth‐ er biomimetic study, conducted by Habibovic *et al.* [26], indicated that a thick and homo‐ genenous crystalline hyroxyapatite coating was deposited on all pores and resembled bone minerals.

An evaporation-based biomimetic coating was introduced by Duan *et al*. [27]. In their study, a supersaturated calcium phosphate was prepared by mixing NaCl, CaCl2, HCl, NH4H2PO4, tri(hydroxymethyl)aminomethane (Tris), and distilled water, which they termed the acceler‐ ated calcification solution (ACS). Calcium phosphate crystallites formed on the surface on dipping the samples into the ACS. The main component in the coating was octa-calciumphosphate (OCP) and apatite was observed after soaking in SBF. The advantages of this method include (i) no surface etching is required, (ii) high supersaturations of the coating chemistry can be achieved, and (iii) tight control of the solutions is achieved [27].

#### **5.4. Thermal spray**

they reported that the coating presented good adhesion to the underlying substrate. The electrodeposition of CaP showed that calcium phosphate enhances the adherence of cells.

Adamek *et al.* [24] succeeded in producing an HA coating on porous Ti6Al4V using electro‐ deposition. The intermediate layer between the bone and metallic implant was rough and porous. Large pores and nanolamellae were present within the HA layer. The flexibility of the electrodeposition technique for coating solid and porous metallic implants has acquired increasing interest due to this ability to enhance the bioactivity of bone implant materials.

There are two major steps involved in the biomimetic technique. The first step is to conduct a pre-treatment of the implant material surface to create a layer functional group that can induce formation of an effective apatite layer. Several studies have revealed that an apatite layer has not been formed on materials without any treatment prior to immersion [25]. Pre‐ liminary treatment includes, for example, hydrothermal, sol-gel, alkali heat treatment and micro-arc. The second step is to immerse the biomaterials into a simulated body fluid (SBF). In this step, the bone apatite layer is formed on the biomaterial's surface. The high apatite forming ability of titanium arises from the formation of a hydrated titanate surface layer during chemical treatment. The advantages of the biomimetic process include (i) flexibility in the control of the chemical composition and thickness of the coating, (ii) the formation of relatively homogenous bioactive bonelike apatite coatings, (iii) a lower processing tempera‐

Wang *et al.* [25] used a modified biomimetic approach to improve the biocompatibility of porous titanium alloy scaffolds. In their experiment, porous Ti10Nb10Zr underwent an alka‐ li heat treatment prior to soaking in SBF. Two NaOH concentrations of 5 M and 0.5 M were used, and the samples were soaked for 1 week. The surface morphologies of porous TiNbZr after alkali soaking and heat treatment revealed a nanofiber layer, that consisted of sodium titanate. Parameters that influenced the morphology and thickness of the sodium titanate

Calcium phosphate was successfully deposited on the surface of the porous TiNbZr. The calcium phosphate layer was uniform and homogenously spread onto the surface. Anoth‐ er biomimetic study, conducted by Habibovic *et al.* [26], indicated that a thick and homo‐ genenous crystalline hyroxyapatite coating was deposited on all pores and resembled

An evaporation-based biomimetic coating was introduced by Duan *et al*. [27]. In their study, a supersaturated calcium phosphate was prepared by mixing NaCl, CaCl2, HCl, NH4H2PO4, tri(hydroxymethyl)aminomethane (Tris), and distilled water, which they termed the acceler‐ ated calcification solution (ACS). Calcium phosphate crystallites formed on the surface on dipping the samples into the ACS. The main component in the coating was octa-calciumphosphate (OCP) and apatite was observed after soaking in SBF. The advantages of this method include (i) no surface etching is required, (ii) high supersaturations of the coating

chemistry can be achieved, and (iii) tight control of the solutions is achieved [27].

**5.3. Biomimetic creation of surfaces**

28 Titanium Alloys - Advances in Properties Control

ture, and (iv) the ability to coat 3D geometries.

were reaction temperature and NaOH concentration.

bone minerals.

The thermal spray technique is a well-established and versatile technique that can be ap‐ plied for a wide variety of coating materials, *i*.*e*., metallic, non-metallic, ceramic, and poly‐ meric. Thermal spray coated medical implants, such as orthopaedic and dental prostheses, have been commercially used. Thermal spray offers several advantages, such as the ability to coat low and high melting materials, a high deposition rate, and flexibility in coating 3D shape components. It is also cost effective [28]. Despite these advantages, some problems have been revealed after long term implantations using thermal spray coatings, such as de‐ lamination, resorption, biodegradation of the thick coating and mechanical instability [29]. Thus, improving the adhesion strength of thermal sprayed coatings is a major concern for bone or dental applications.

There are several types of thermal spraying; for example, plasma spray, flame spray, and cold spray [29]. Plasma spray is commonly applied to produce thick coatings for metallic corrosion protection. It is also flexible, due to its ability to coat different substrates. During plasma spraying, the precursor is atomised and injected into plasma jet, then accelerated to‐ wards the substrate with the aid of an inert carrier gas [30]. There are many parameters that must be controlled to produce a high quality coating.

Flame spray uses a combustion flame to melt the solid precursor. There is, additionally, an‐ other type of flame spray termed as high velocity oxygen fuel (HVOF). This technology is favourable due to its high spray velocity and the formation of a strong bond coating [30-32].

The thermal spray technique has been widely employed for HA coatings. The surface morphology of HA coatings obtained with various parameters of stand-off distance (SOD) and power are presented in Figure 2 (a)-(d). Sun *et al.* [28] reported that when the spray power increased, the crystallinity of HA decreased and the amorphous phase became more obvious. The effect of SOD indicated an inverse correlation with deposition efficien‐ cy. Several parameters that influence the deposition of HA are SOD, spray power, feed‐ stock particle size and velocity.

Cold spray is a new member of the thermal spray family. This technique uses small particles of 1-50 µm. A supersonic jet of compressed gas is used to accelerate the particles. The ad‐ vantage of using this technique is the ability to produce dense coatings and maintain the material chemistry and phase composition of the feedstock. Noppakun *et al.* [33] have ap‐ plied cold spray technique to deposit HA-Ag/poly-esther-ether-ketone on glass slides. This study reported that cold spray was able to retain and elicit a coating functionality that was the same as the starting materials.

#### **5.5. Physical vapor deposition**

Physical vapor deposition (PVD) is a deposition method where materials are evaporated or sputtered, transferred and deposited onto the substrate surface. This physical process in‐ cludes thermal evaporation or plasma-induced ion bombardment onto the sputtering target. A condensation or reaction of the coating materials then takes place on the substrate surface to form coatings. Variants of the PVD process include evaporation, ion plating, pulsed laser dep‐ osition and sputtering. The beneficial features of PVD are high coating density, high bio-adhe‐ sion strength, formation of multi-component layers, and low substrate temperature [34].

sectors. There are several sputtering methods, such as DC glow discharge, radio frequency

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The simplest model for sputtering is the diode plasma, which consists of a pair of planar electrodes, an anode and a cathode, inside a vacuum system [37]. The sputtering target is mounted on the cathode. Application of the appropriate potential difference between the cathode and anode will ionize argon gas and create a plasma discharge. The argon ions will then be attracted and accelerated toward the sputtering target. Such ion bombard‐ ment on the target will displace some of the target atoms. This results in electron emis‐ sion that will subsequently collide with gas atoms to form more ions that sustain the discharge [37]. Ion beam sputtering has disadvantages, such as a high capital investment cost (approximately one million dollars per machine), low deposition rates and a relative‐ ly small capacity per chamber batch [38]. Another type of sputtering employs radio fre‐

Magnetron sputtering is one option to overcome the problems such as delamination and low bond strength that may arise with plasma spray methods. Magnetron sputtering ena‐ bles lower pressures to be used, because a magnetic field allows trapping of the secondary electrons near the target. This induces more collisions with neutral gases and increases plas‐ ma ionisation. Figure 3 is a diagram of the magnetron sputtering mechanism. RF magnetron sputtering is an improved ion-sputtering method. It has also been noted that sputtered films

possess higher adhesion to the substrate compared to the evaporation method.

(RF), ion beam sputtering (IBS), and reactive sputtering [36].

**Figure 3.** Schematic diagram of the sputtering mechanism

quency (RF) diodes that operate at high frequency.

**Figure 2.** Surface morphology of HA coatings obtained by thermal spray method (a) 27.5 kW at 80 mm SOD, (b) 27.5 kW at 160 mm SOD, (c) 42 kW at 80 mm SOD, and (d) 42 kW at 160 mm SOD

Evaporation involves the thermal phase change from solid to vapor under vacuum condi‐ tions, in which evaporated atoms of a solid precursor placed in an open crucible can travel directly and condense onto the surface of a substrate [35]. A vacuum environment is used to minimize contamination [36]. Han *et al.* [37] have created an HA coating using electron beam evaporation and then incorporated silver by immersion into AgNO3 solution. One ad‐ vantage of this method is an improved bond strength between the coating and substrate. The ratio of Ca/P in the HA coating was 1.62 with a bond strength of 64.8 MPa, which was significantly higher than a plasma sprayed bond strength of 5.3 MPa [37].

Sputtering involves a process of ejecting neutral atoms from a target surface using energetic particle bombardment. The energetic particles used in the sputtering process are argon ions, which can be easily accelerated towards the cathode by means of an applied electric poten‐ tial, hence bombarding the target, and ejecting neutral atoms from the target. These ejected atoms are then transferred and condense to the substrate to form a coating. Sputtering has been used in many applications such as the semiconductor, photovoltaic and automotive sectors. There are several sputtering methods, such as DC glow discharge, radio frequency (RF), ion beam sputtering (IBS), and reactive sputtering [36].

**Figure 3.** Schematic diagram of the sputtering mechanism

osition and sputtering. The beneficial features of PVD are high coating density, high bio-adhe‐ sion strength, formation of multi-component layers, and low substrate temperature [34].

30 Titanium Alloys - Advances in Properties Control

**Figure 2.** Surface morphology of HA coatings obtained by thermal spray method (a) 27.5 kW at 80 mm SOD, (b) 27.5

Evaporation involves the thermal phase change from solid to vapor under vacuum condi‐ tions, in which evaporated atoms of a solid precursor placed in an open crucible can travel directly and condense onto the surface of a substrate [35]. A vacuum environment is used to minimize contamination [36]. Han *et al.* [37] have created an HA coating using electron beam evaporation and then incorporated silver by immersion into AgNO3 solution. One ad‐ vantage of this method is an improved bond strength between the coating and substrate. The ratio of Ca/P in the HA coating was 1.62 with a bond strength of 64.8 MPa, which was

Sputtering involves a process of ejecting neutral atoms from a target surface using energetic particle bombardment. The energetic particles used in the sputtering process are argon ions, which can be easily accelerated towards the cathode by means of an applied electric poten‐ tial, hence bombarding the target, and ejecting neutral atoms from the target. These ejected atoms are then transferred and condense to the substrate to form a coating. Sputtering has been used in many applications such as the semiconductor, photovoltaic and automotive

kW at 160 mm SOD, (c) 42 kW at 80 mm SOD, and (d) 42 kW at 160 mm SOD

significantly higher than a plasma sprayed bond strength of 5.3 MPa [37].

The simplest model for sputtering is the diode plasma, which consists of a pair of planar electrodes, an anode and a cathode, inside a vacuum system [37]. The sputtering target is mounted on the cathode. Application of the appropriate potential difference between the cathode and anode will ionize argon gas and create a plasma discharge. The argon ions will then be attracted and accelerated toward the sputtering target. Such ion bombard‐ ment on the target will displace some of the target atoms. This results in electron emis‐ sion that will subsequently collide with gas atoms to form more ions that sustain the discharge [37]. Ion beam sputtering has disadvantages, such as a high capital investment cost (approximately one million dollars per machine), low deposition rates and a relative‐ ly small capacity per chamber batch [38]. Another type of sputtering employs radio fre‐ quency (RF) diodes that operate at high frequency.

Magnetron sputtering is one option to overcome the problems such as delamination and low bond strength that may arise with plasma spray methods. Magnetron sputtering ena‐ bles lower pressures to be used, because a magnetic field allows trapping of the secondary electrons near the target. This induces more collisions with neutral gases and increases plas‐ ma ionisation. Figure 3 is a diagram of the magnetron sputtering mechanism. RF magnetron sputtering is an improved ion-sputtering method. It has also been noted that sputtered films possess higher adhesion to the substrate compared to the evaporation method.

A summary of the characteristics of the various coating techniques for calcium phosphate is presented in Table 2. Each technique has its own benefits and drawbacks. However, sputter‐ ing is a promising method due to its ability to produce dense and thin coatings, as well as provide good bond strength [39-41].

was higher than coatings without a sublayer. Nieh *et al.* [48] used titanium as a pre-coat on

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Layered materials have previously been demonstrated to improve bonding between dissim‐ ilar materials. According to Ding [45], the top layer provides an excellent interaction with the surrounding tissue and promotes bone healing. A functionally graded coating (FGC) is an alternative method to enhance coating adhesion strength. Ozeki *et al.* [49] prepared an FGC of HA/Ti onto a metallic substrate. The coating thickness was 1 µm and consisted of 5 layers. The configuration of FGC was designed so that the HA was more dense near the sur‐ face, whilst the Ti was more dense near the substrate. The bonding strength using the FGC configuration was higher than using only HA, *i*.*e*., 15.2 MPa and 8 MPa for the FGC and

*Elastic properties.* Snyders *et al.* [50] manufactured HA via RF sputtering and revealed that the chemical composition influenced the elastic properties. As the Ca/P ratio decreased, the elastic modulus also decreased due to the insertion of Ca vacancies in the HA lattices.

The biological behaviour of biomaterials has been a fundamental criterion for successful candidate implant materials, along with their mechanical properties. The surface properties of a biomaterial play a significant role in the cell response. Thus, surface modification is an established strategy that has been used for biomedical applications due to its ability to en‐ hance bioactivity. High cell density enhances bone formation. The cell adhesion behaviour and proliferation are influenced by several factors, such as pore size, porosity, and surface

Thian *et al.* [45] carried out an *in vitro* test using a human osteoblast (HOB) cell model for a silicon incorporated hydroxyapatite (Si-HA) coating on titanium. The sample demon‐ strated an increase in metabolic activity compared to mono-HA coatings. Sputtered HA and Si coatings exhibited good differentiation of osteogenic cells and good biocompatibili‐ ty. It was noted that the biological response was influenced by the crystallinity of the HA coatings. Sputtered composite coatings of HA with other compounds may provide addi‐ tional advantages for implant performance. For instance, Chen *et al*. [52] incorporated sil‐ ver into HA, conducted a cytotoxic and antibacterial test, and reported that the silver had an antibacterial effect since the bacterial attachment was reduced compared to coatings

Tin and niobium were chosen as alloying elements because both metals are biocompatible and non-cytotoxic. The titanium alloy composition was designed using the molecular orbital DV-Xα method [53]. The calculation of the nominal composition of the alloys was based on

Ti6Al4V and found strong bonding between the Ti layer and the HA coating.

*5.5.2. Biological performance of sputtered hydroxyapatite coatings*

pure HA, respectively.

composition [51].

that did not contain silver.

**6. Experimental methods**

**6.1. Design and preparation of titanium alloys**


**Table 2.** Summary of various techniques for calcium phosphate coatings

#### *5.5.1. Properties of sputtered hydroxyapatite coatings*

*Coating thickness.* The HA coating thickness varies. Molagic [42] succeeded in producing HA/ZrO2 coatings with an average thickness of 3.2 µm. Hong *et al.* [43] manufactured a 500 nm thick coating of crystalline HA using magnetron sputtering. Ding [44] sputter de‐ posited HA/Ti coatings with a film thickness of 3-7 µm onto a titanium substrate. Thian *et al.* [45] succeeded in incorporating silicon in hydroxyapatite (Si-HA) using magnetron sputtering and discovered its potential use as a bio-coating. The Si-HA film thickness was up to 700 nm.

*Bond strength.* An *in vitro* and *in vivo* experiment on coatings using the sputtering technique revealed coating detachment problems. Cooley *et al.* [46] reported that HA coatings were re‐ moved after 3 weeks of implantation. A bond layer coating was suggested to overcome this weak adhesion at the interface and subsequent delamination. Ievlev *et al.* [47] measured the adhesion strength of HA coatings with a sublayer and revealed that the adhesion strength was higher than coatings without a sublayer. Nieh *et al.* [48] used titanium as a pre-coat on Ti6Al4V and found strong bonding between the Ti layer and the HA coating.

Layered materials have previously been demonstrated to improve bonding between dissim‐ ilar materials. According to Ding [45], the top layer provides an excellent interaction with the surrounding tissue and promotes bone healing. A functionally graded coating (FGC) is an alternative method to enhance coating adhesion strength. Ozeki *et al.* [49] prepared an FGC of HA/Ti onto a metallic substrate. The coating thickness was 1 µm and consisted of 5 layers. The configuration of FGC was designed so that the HA was more dense near the sur‐ face, whilst the Ti was more dense near the substrate. The bonding strength using the FGC configuration was higher than using only HA, *i*.*e*., 15.2 MPa and 8 MPa for the FGC and pure HA, respectively.

*Elastic properties.* Snyders *et al.* [50] manufactured HA via RF sputtering and revealed that the chemical composition influenced the elastic properties. As the Ca/P ratio decreased, the elastic modulus also decreased due to the insertion of Ca vacancies in the HA lattices.

#### *5.5.2. Biological performance of sputtered hydroxyapatite coatings*

The biological behaviour of biomaterials has been a fundamental criterion for successful candidate implant materials, along with their mechanical properties. The surface properties of a biomaterial play a significant role in the cell response. Thus, surface modification is an established strategy that has been used for biomedical applications due to its ability to en‐ hance bioactivity. High cell density enhances bone formation. The cell adhesion behaviour and proliferation are influenced by several factors, such as pore size, porosity, and surface composition [51].

Thian *et al.* [45] carried out an *in vitro* test using a human osteoblast (HOB) cell model for a silicon incorporated hydroxyapatite (Si-HA) coating on titanium. The sample demon‐ strated an increase in metabolic activity compared to mono-HA coatings. Sputtered HA and Si coatings exhibited good differentiation of osteogenic cells and good biocompatibili‐ ty. It was noted that the biological response was influenced by the crystallinity of the HA coatings. Sputtered composite coatings of HA with other compounds may provide addi‐ tional advantages for implant performance. For instance, Chen *et al*. [52] incorporated sil‐ ver into HA, conducted a cytotoxic and antibacterial test, and reported that the silver had an antibacterial effect since the bacterial attachment was reduced compared to coatings that did not contain silver.
