A short review: hydroxyapatite coatings for metallic implants

ABSTRACT

The number of orthopedic surgeries will continue increasing over the next decades. The most common causes of orthopedic implant failure after implantation are infections, inflammatory response, corrosion, mismatch in elastic modulus, stress shielding and excessive wear. To address the problems associated with implants, different modifications related to geometry, materials and surface have been developed. Among the different methods, coating is an effective method to improve the performance of implant materials. Hydroxyapatite is the best option to coat metallic implants and it has been used in clinical practice for more than 30 years. In this article, a comprehensive review of recent studies has been carried out to summarize the impact of hydroxyapatite on metallic implants. The antibacterial characteristics, biodegradability, biocompatibility, corrosion behavior and mechanical properties for performance evaluation are briefly summarized. Different effective coating techniques and additives have been summarized. The results are useful to produce the coating with optimized properties.

KEYWORDS:

• Orthopedic implants
• coatings
• hydroxyapatite
• biocompatibility

Introduction

Today, every country in the world is experiencing growth in both the size and the proportion of older persons in the population. By 2030, people aged 60 and over will increase from 1 billion in 2020 to 1.4 billion and they will represent 16% of the world’s population. By 2050, the world’s population of people aged 60 years and older will double (2.1 billion). Between 2020 and 2050, the number of persons aged 80 years or older is expected to triple to reach 426 million [1,2]. The aging of the population and the fact that patients in their middle years are increasingly choosing for orthopedic implants due to lifestyle-related concerns such as early burnout and lack of exercise are increasing the number of orthopedic surgeries. In the United States, the size of the orthopedic implant market in 2021 was $45.30 billion and is estimated to reach $64.18 billion by 2028, with joint reconstruction representing the largest share among other product types, such as orthobiologics, spinal, trauma and dental implants [3]. These implants are required to be inert to living tissues, not carcinogenic, not toxic, not dense, strong, elastic, ductile, tough, corrosion resistant and must promote osteointegration [4–6]. Many materials including metals, alloys, ceramics, polymers and composites are used in orthopedic applications [7]. Among these materials, metallic materials have shown better mechanical properties (strength, elastic modulus ductility and toughness). The most common metals used in orthopedic applications are stainless steel, titanium alloys and cobalt-chromium alloys due to their cost-effectiveness, excellent durability and favorable mechanical properties; on the other hand, biodegradable metals, such as magnesium, iron and zinc are used for temporary applications [8–13]. Although metallic implants, especially titanium alloys have great mechanical properties, they do not show good wear and corrosion resistance, as well as an ability to support osseointegration over the long term [14]. For this reason, it is necessary to develop novel orthopedic implants for bone and joint replacement or to support damaged bones [15–17]. One approach to overcome this issue is to coat the surface of the metallic implants with functional materials that alter surface morphology and topography, energy, charge and chemistry [18]. A good coating should have sufficient adhesion strength, high hardness, excellent osseointegration, good osteoconduction, low number of cracks and no inclusions. Calcium phosphates are perfect candidates for this purpose, since 65% of the mineral phase of bone is composed of Ca₅(PO₄)₃(OH) (hydroxyapatite [HA]), a calcium phosphate; nevertheless, it is important to consider that only some calcium phosphates are useful for this purpose, for its solubility increases when its Ca/P ratio is lowered. Crystalline HA, which has a hexagonal crystalline structure, spatial group P63/m, cell parameters a = b = 9.42 Å and c = 6.88 Å, is considered the best option to coat metallic implants and has been used in clinical practice for 30 years [3,19,20]. Figure 1 shows hydroxyapatite’s crystalline structure.

Figure 1. Crystalline structure of hydroxyapatite [21].

Table 1 shows some mechanical properties of HA.

Table 1. Mechanical properties of HA [22].

Property	Value
Theoretical density	3.156 g/cm^3
Hardness	500–800 HV
Tensile strength	40–100 MPa
Bend strength	20–80 MPa
Compressive strength	100–900 MPa
Fracture toughness	Approx. 1 MPa m^0.5
Young’s modulus	70–120 GPa

Table 2 shows ISO and ASTM requirements for HA coatings. Different techniques such as electrophoretic deposition, ion-beam sputter coating, dip coating, plasma spraying, etc. are used to coat metallic implants with HA [14]. The aim of this article is to review and compare the properties of synthesized HA coatings on metallic substrates obtained with different deposition techniques.

Table 2. Required ISO and ASTM specifications for HA coatings [23].

Property	Value
Thickness	50–150 µm
Crystallinity	Over 62%
Purity	Over 95%
Ca/P Ratio	1.67–1.76
Density	2.98 g/cm^3
Heavy metal concentration	Less than 50 ppm
Normal tension	Over 50.8 MPa
Shear stress	Over 22 MPa

Techniques to deposit HA coatings on implant materials

RF sputtering

RF Sputtering runs an energetic wave which ionize atoms of an inert gas in a vacuum chamber. The target material (cathode) is bombarded by these high-energy ions, removing atoms that are deposited into the substrate (anode). After some time, the positive ions start to accumulate on the surface of the target, giving it a positive charge. Later, this charge can build up and provoke a secession of atoms beings discharged for coating. To avoid this issue, RF Sputtering alternates the electric potential. In this way, on the positive cycle, electrons are attracted to the target material giving it a negative bias, while in the negative cycle ions are attracted to the target material giving it a positive bias to continue the sputtering. RF sputtering offers different advantages depending on the material obtained. RF plasma diffuses through all the chamber rather than just near the target material. RF plasma can be sustained throughout the chamber with low pressures, avoiding collisions between ions, which makes the deposition more efficient. As the target material is not allowed to build up charge, arcing is reduced. RF sputtering reduces the creation of a ‘racetrack erosion’ on the surface of the target material. Contrary to DC sputtering, in this technique, there is no disappearing anode effect when the substrate becomes insulated and acquires a charge. Although RF sputtering offers different advantages depending on the type of material to be coated, there are several disadvantages that must be considered. RF sputtering uses radio waves instead of DC current; for this reason, deposition rates are slower and require higher voltages in RF sputtering, which may result in overheating and the use of more expensive power supplies. This disadvantage makes RF sputtering being more suitable for small coatings. Another disadvantage of RF sputtering is the decrease in deposition rates due to the lack of available electrons being trapped above the target material for gas ionization. RF sputtering may be used to coat different type of materials; however, it has become the technique of choice to deposit dielectric (non-conducting) coatings that can take on a polarized charge [24,25]. Figure 2 shows a schematic of this technique.

Figure 2. Diagram of the RF Sputtering Process [24].

Lemoine et al. [26] deposited HA films on AZ31 Mg alloy substrates using RF sputtering and analyzed their morphology, composition, structure, roughness and nano-mechanical properties. EDX and XRD analysis showed that the coatings were amorphous and Ca-deficient and that after thermal annealing they had a near-stochiometric composition. Table 3 shows EDX analysis of the samples. Tapping AFM microscopy (TAFM) analysis showed that despite being rougher, the samples displayed better wear resistance. Nanoindentation analysis showed that annealing increased the intrinsic hardness (from 2.7 GPa to 9.4 GPa) and strain at break (from 0.043 to 0.079). Figure 3 shows the Young’s modulus (E) and hardness (H) values obtained at a 200 nm depth.

Figure 3. Bar charts of E and H values of the samples. H100 (sample deposited during 100 h), H100-A1 (sample deposited during 100 h and treated once to ∼600°C for 15 s) and H100-A2 (sample deposited during 100 h and treated twice to ∼600°C for 15 s) [26].

Table 3. Mechanical properties of uncoated FeMoTaTiZr alloy, doped and undoped hydroxyapatite (HAP) coatings deposited on FeMoTaTiZr alloy [27].

Sample	Ra (mm)	Hf (GPa)	Ef (GPa)
Uncoated FeMoTaTiZr	85.1 ± 21.1	10.2 ± 0.5	129.3 ± 7.7
FeMoTaTiZr coated by HAP	122.3 ± 37.7	5.5 ± 0.1	111.9 ± 6.9
FeMoTaTiZr coated by HAP + Zn	207.8 ± 61.5	6.8 ± 0.1	108.5 ± 14.8

Codescu et al. [27] obtained HA coatings undoped and doped with Zn on FeMoTaTiZr substrates using RF sputtering and analyzed their morphology, elemental composition, phase composition, mechanical properties, roughness and corrosion resistance. EDS analysis confirmed the presence of Ca and P in the case of all coatings, while the Ca/P ratio was 1.69 for the undoped HA and 1.96 for the Zn-doped HA. XRD and ATR-FTIR analysis showed the presence of calcium phosphate phases. The roughness of the HA-coated substrates increased 1.4 times, whereas it increased 2.4 for the case of HA + Zn-coated substrates, compared to the uncoated substrates. The nanoindentation results showed that the Zn-based HA coatings can be used to decrease the elastic modulus of biocompatible alloys. Table 3 shows the results obtained for the roughness, hardness and Young modulus measurements. Electrochemical analysis showed that the undoped HA coatings had good corrosion resistance, and that Zn-doped HA coatings had a high dissolution rate in Fetal Bovine Serum (FBS), suggesting that Zn-doped HA coatings could serve better as a biodegradable material. Table 3 shows the mechanical properties of uncoated FeMoTaTiZr alloy, doped and undoped hydroxyapatite (HAP) coatings deposited on FeMoTaTiZr alloy.

Ivanova et al. [28] coated titanium substrates with HA with a,b-plain and c-plain preferential orientation using RF-magnetron sputtering to analyze the influence of microstructural features on mechanical behavior of the coating. Nanoindentation results showed that the (002) and (300) textured coatings have a nanohardness of 4.7 ± 2.0 GPa and 4.4 ± 2.2 GPa, respectively; on the other hand, the elastic modulus of the HA film with (300) and (002) preferential orientation was 75 ± 40 GPa and 103 ± 40 GPa, respectively. The HA coating with (300) texture showed an elastic recovery of 92%, whereas the (002) textured film had an elastic recovery of 69%. Figure 4 shows the nanohardness H and Young’s Modulus E as a function of the indentation depth h for HA coatings.

Figure 4. The nanohardness H (a,b) and Young’s modulus E (c,d) values plotted as a function of the indentation depth h for HA coatings deposited via RF-magnetron sputtering textured along c-axis (a,c) and a-axis (b,d). Error bars are the standard deviations of those measurements [28].

Microwave-assisted synthesis

Microwave chemistry is based on the use of dielectric heating effects to efficiently heat materials. There are two main mechanisms for dielectric heating: dipolar polarization and ionic conduction. In dipolar polarization, dipoles are exposed to an oscillating microwave field and then align to it, resulting in rotation that results in friction and ultimately in heat energy. In ionic conduction, dissolved charged particles oscillate under the influence of microwave irradiation. These oscillations cause collisions with neighboring atoms or molecules which end up creating heat energy. The heating features of a particular material are determined by their ability to convert electromagnetic energy into heat when they are under the influence of microwave irradiation. This ability is measured by the so-called loss tangent, tan δ. Compared to conventional heating, microwave irradiation heating results in a more energy efficient internal heating, for it couples microwave energy with dipoles and/or ions that are present in the reaction mixture. Microwaves pass through the microwave-transparent vessel wall and heat the reaction mixture by direct interaction with the molecules; therefore, inverted temperature gradients are obtained as compared to the conventionally heated system, as is illustrated in Figure 5. Moreover, microwave heating occurs at a faster rate than conventional heating, so the formation of byproducts is suppressed in the first one. Microwave accelerated reactions are carried out in domestic microwave oven or especially designed microwave equipment. This equipment has a fixed frequency of 2.45 GHz and work in between 500 and 1500 W power. Microwaves are generated by magnetron and the temperature is maintained by turning this on and off. All types of solvents can be used in microwave synthesis, but polar solvents are more suitable since they are able to absorb microwave energy. Even if non-polar solvents are used, the reaction mixture contains enough polar or ionic substance, which can absorb microwave irradiation and generate heat to complete the reactions. There are two types of microwave-assisted synthesis: microwave-assisted synthesis using solvents and microwave-assisted synthesis without solvents [29,30].

Figure 5. Illustration of heat introduction and temperature distribution in a reaction mixture for (a) conventional heating and (b) microwave heating [29].

Panda et al. [31] deposited HA coatings reinforced with strontium and niobium on Ti6Al4V alloy sheets using microwave (MW) irradiation technique and analyzed their physical properties and biocompatibility. SEM, EDS, XRD and FTIR analysis showed that the coatings were uniform, semi-crystalline and were made of HA. Indentation and roughness analysis showed that the coated samples increased microhardness and surface roughness with respect to uncoated samples. Coated samples also experienced higher wettability and more protein adsorption. Electrochemical corrosion evaluation showed an increase in corrosion potential and a decline in corrosion current density, which suggest an anticorrosive behavior. These samples also exhibited improved hemocompatibility and bacteriostatic properties. Finally, cell viability and confocal microscopy analysis showed an increment in cell attachment on the surface. Figure 6 shows the results of cell viability.

Figure 6. Percentage cell viability of uncoated and coated samples. Ctrl (bare Ti6Al4V), A (cleaned Ti6Al4V), SN0, SN1 and SN2 (HA coatings with strontium and niobium with different molar ratios) [31].

Panda et al. [32] obtained niobium-doped hydroxyapatite (Nb-HAP) coatings on grade-5 titanium (Ti6Al4V) substrates by microwave irradiation. The coated samples contained calcium-deficient hydroxyapatite (HAP) with niobium (Nb). The apatite layer on the substrates adsorbs more protein and enhanced the cell survival by more than 30%. The coatings appear to be denser, thicker and crystalline after doping. Samples became more hydrophilic and increased corrosion resistance after doping. The presence of Nb in the coating in the coating created a biofilm property against both gram −ve and gram +ve bacteria strains. It was also observed an improved hemocompatibility towards goat erythrocytes and cytocompatibility towards human keratinocyte cells after coating. The microwave-assisted synthesis of Nb-doped HAP coating improved the physical properties of the substrate and increase the biocompatibility of bone implant. For this reason, the microwave irradiation technique appears as a fast and economic procedure for bioactive HAP coating deposition. Figure 7 shows the results of cell viability.

Figure 7. Percentage viability of cells (cell viability test performed using MTT assay) [32].

Shen et al. [33] coated HA on magnesium alloys in an aqueous solution under microwave irradiation at 100°C in 10 min. The results showed that microwave irradiation promoted the growth of hydroxyapatite nuclei into rod-like crystal, and the thickness of the coatings was ∼9.9 μm, larger that the one obtained by water-bath heating, which was ∼6.3 μm. The coating obtained by microwave-assisted deposition offered more corrosion resistance for the magnesium alloy compared with the coating obtained by water-bath heating. The corrosion current density of the coated substrates was reduced by 100-fold as compared to the bare substrates. Figure 8 shows the polarization curves.

Figure 8. Polarization curves [33].

Plasma spray

Plasma spray consists of an anode (cooper), a cathode (tungsten), a power supply, a gas (argon, nitrogen or a mixture of them combined with hydrogen and helium), cooling water, a carrier gas injection nozzle and a powdered source compound. The gas flows between the cathode and the anode, which is used as an injector and experiences a high frequency discharge generated by the application of an electric current. In this way, a plasma with free electrons, ionized atoms, some neutral atoms and non-dissociated atomic molecules is produced. The speed and temperature of this gas are determined by the applied voltage between the anode and cathode, its density, its flow and the electric power. Voltage and current depend on the electrode’s design, the flow and the composition of the gas. In most cases, the speed of the plasma is subsonic; however, in some cases, a supersonic speed can be reached in injectors with the proper critical angle. The temperature of the plasma nucleus can exceed 30,000°C. Once the plasma leaves the injector, it carries the powdered source compound. The most important parameters when the powder reaches the substrate are its temperature, speed and reaction with the plasma. The speed reached by the powder depends on the plasma flow and its trajectory, while its temperature depends on transport time, plasma’s temperature and plasma’s composition. It is commonly mentioned that any material that experiences melting without decomposition can be used as a coating; nevertheless, the fluidity level and the speed of the powder particles should be optimum, so that take the shape of the substrate’s surface. This technique offers a high deposit rate, but it is expensive and generally the powder reacts with the oxygen and nitrogen contained in the plasma; for this reason, it is sometimes necessary to modify the plasma [34]. Figure 9 shows a schematic of this technique.

Figure 9. Schematic of the Plasma Spray deposition technique [35].

Singh et al. [36] deposited pure HA and HA + 7wt.% Aloxite coatings on pure titanium substrates using plasma spray and analyzed their surface morphology and composition before and after heat treatments. XRD analysis revealed that HA with some amorphous phases were obtained before heat treatment, crystalline HA with some CaO traces were obtained after heat treatment at 750°C and pure crystalline HA was obtained after heat treatment at 950°C. Moreover, no substrate element was identified in the X-ray mapping of coatings. Figure 10 shows the XRD plot of HA-R7 coated and heat-treated pure titanium specimens. After heat treatment of coatings, ultrafine particles were observed on the surface of the coatings. These ultrafine particles have the capability to bond with bone and promote bone growth on their surfaces.

Figure 10. XRD pattern of HA – R7 coating pure titanium: (a) as coated; after heat treatment for 2 h in air at (b) HT-750 and (c) HT-950 [36].

Hussain et al. [37] obtained HA coatings on titanium substrates using plasma spray. The microstructure and phases were analyzed with a field-emission scanning electron microscope (FESEM) and an X-ray diffractometer (XRD). The surface roughness, microhardness, porosity, adhesion strength and wear resistance were investigated with a stylus profilometer, a Vickers microhardness tester, image analysis technique, scratch tester and ball-on-dic tribometer, respectively. The average surface roughness (Ra) and porosity of the coating decreased when the synthesis temperature increased. It was suggested that the low roughness and porosity is due to a high degree of melting of the powder particles. The samples with the highest synthesis temperature showed the highest microhardness and wear-resistance due to its maximum crystallinity among all the samples obtained. Figure 11 shows the average coating roughness as a function of synthesis temperature.

Figure 11. Variation of coating roughness of different plasma sprayed HA coatings [37].

Rattan et al. [38] coated titanium substrates with HA and HA reinforced with 15 wt-% aluminum oxide (Al₂O₃). SEM, EDX and XRD were used to investigate the coated samples. A Tribometer (Ball on a disc assembly) was used to perform wear tests. The wear resistance in the HA reinforced with 15 wt-% aluminum oxide (Al₂O₃) coatings (wear rate = 4.9 × 10⁻³ mm³/N-m) was better than that of HA coatings (wear rate = 8.4 × 10⁻³ mm³/N-m). These results show that adding Al₂O₃ to the coatings considerably improves their tribological properties. Figure 12 shows friction coefficient curves for the samples.

Figure 12. Friction coefficient curves of HA and HA + 15 wt-% Al₂O₃ coatings [38].

Laser-induced single-step coating

Laser-induced single-step coating uses the effect of hydrothermal synthesis on a laser-irradiated solid (substrate)/liquid (precursor solution) interface. Deposition from the liquid state has advantages over deposition from other states of matter in terms of relative ease of handling, simple setup and recyclability. Moreover, with this technique, it is possible to both synthesize and coat on the surface that is in contact with the laser. For this reason, several coating layers of the desired materials can be obtained depending on the type of immersion solution. If the focal length of the laser is not automatically adjusted, then it may be difficult to coat a complicated structural surface [39–41]. Figure 13 shows a schematic of this technique.

Figure 13. Schematic of the Laser-Induced Single-Step Coating deposition technique. MOPA Fiber laser (Master oscillator power amplifier fiber laser), F-theta lens (Lenses designed to focus a laser beam onto a planar image plane) [40].

Oyane et al. [42] deposited HA coatings on titanium substrates using induced-laser deposition. Precipitation occurred due to the effect of laser surface modification and ambient heating. Moreover, there was immobilization of various osteogenic substances (zin and fibronectin components) on the titanium surface with HA by adding them to the HA solution. This process is expected to be an easy and effective way to provide a titanium surface with osteoconductivity. Figure 14 shows scanning electron microscopy images of untreated titanium substrate surface and that of a coating.

Figure 14. Scanning electron microscopy (SEM) images of the untreated titanium substrate surface and that of a coating. LAB (laser single-step coating) [42].

Park et al. [39] obtained HA coatings on Mg substrates using laser-induced single-step coating as a deposition technique and solutions of calcium and phosphoric acid as precursors. It was found that corrosion decreased, and biocompatibility increased when Mg substrates were coated. Figure 9 shows cell adhesion results. This surface coating technique offers a simple process that doesn’t need chemical ligands, and so overcomes a potential obstacle in its clinical use. Figure 15 shows the cell adhesion area for the samples obtained.

Figure 15. Analysis of the cell adhesion area by evaluating the image processing. Ctrl (control group), LIH (laser-induced Hap coating). The p-value is a result of a t-test between sample groups [39].

Um et al. [41] coated titanium substrates with HA using laser-induced single-step coating technique. The coating layer obtained exhibited a binding force of 31.7–47.2 N, which is sufficient for medical applications. The resulting HA coatings facilitated the attachment of bone cells. A rapid single-step method for simultaneous synthesis and coating of HA via nanosecond laser surface treatment was developed. Figure 12 shows the attached cell number per unit area for the coatings obtained. Figure 16 shows cell adhesion results.

Figure 16. The schematic diagram for evaluating the cell adhesion on each surface. Ti (Titanium), P10-L50 (HA coating synthesized with 10 W and 50 loops), P10-L100 (HA coating synthesized with 10 W and 100 loops). Data represent mean ± SD, **** indicates statistically significant difference of p < .0001, N.S. indicates no statistically significant difference [41].

Electrostatic spray-assisted vapor deposition

Electrostatic Spray-Assisted Vapor Deposition (ESAVD) is a technique used to deposit thin or thick layers of a coating onto a substrate at ambient temperature. The chemical precursors are sprayed across an electrostatic field into a heated substrate. The electrostatic charge is obtained by conveying the chemical precursors through a special spraying device. The charged particles are attracted to the ground part to be coated. These chemicals then undergo a controlled chemical reaction and are deposited on the substrates. ESAVD is used for many applications including: thermal barrier coatings for jet engine turbine blades, thin layers in the manufacture of flat panel displays and photovoltaic panels, electronic components, biomedical coatings, glass coatings and corrosion protection coatings. This technique has advantages over others in that it doesn’t require any vacuum, electron beam or plasma, so the manufacturing costs are reduced. It uses less power and raw materials, which makes it more environmentally friendly. Also, the use of an electrostatic field allows it to coat complex 3D parts easily [43]. Figure 17 shows a schematic of this technique.

Figure 17. Showing a schematic diagram of electrostatic spray-assisted vapor deposition (ESAVD) method [44].

Müller et al. [45] deposited HA coatings on Ti6Al4V substrates using electrostatic spray-assisted vapor deposition and analyzed their morphology, crystal size, composition and cell viability. XRD and FTIR analysis showed that all coatings contained nanocrystalline HA and supported cell viability. Table 4 shows the deposition conditions for the samples. Figure 18 shows that a higher cellular response was found when low crystalline HA was used.

Figure 18. Cell viability at 24 h of osteoblast-like cells (MG-63 cell-line) for selected HAP coatings. Values represent means ± SD (n = 9). * Indicates p < .01, and ** indicates p < .001 [45].

Table 4. ESD processing parameters, nominal Ca/P ratio and estimated physical–chemical properties of the precursor solution [45].

Sample code	Type of solvent	Ca^2+ (Mm)	p^5+ (Mm)	Nominal Ca/P ratio	t (h)	Q (mL h^−1)	T (°C)	d (mm)	Type	Crystallite size (nm)	Thickness
DF	EtOH	0.6	3	0.2	3	1.5	350	30	Thin dense	19 ± 3	70 nm
HC	EtOH	15	18.75	0.8	5	1	350	30	Coral	30 ± 3	26 μm
LC	EtOH	15	75	0.2	5	1	350	30	Coral	16 ± 4	26 μm
RT	MeOH:H2O (19:1)	15	10	1.5	4.5	1.5	350	15	Reticular	32 ± 4	6 μm

Muller et al. [46] coated HA coatings on Ti6Al4V substrates using electrostatic spray deposition. The microstructure and composition of the coatings resulted from a compromise between the physicochemical properties of the precursor solution and the ESD deposition conditions. Results showed that the Ca/P molar ratio and the nature of phases are strongly dependent on hydrolysis and evaporation of the phosphorus precursor in the ESD process. Optimized deposition conditions are found by using solvents of low-boiling point, for which the coatings deposited above 325°C correspond to a well crystalline, nanostructured and single-phase HAP. Infrared heat treatment of a coating prepared from the one optimized precursor solution promotes nanostructured HA. Figure 19 shows the experimental Ca/P molar ratios of the as-deposited coatings.

Figure 19. Experimental Ca/P molar ratio of as-deposited coatings measured by EDS and ICP versus the nominal ones [46].

Gokcekaya et al. [47] coated Ti substrates with Ag-incorporated HA using electrostatic spraying deposition. The coatings obtained with various deposition times and heat treatment conditions were analyzed by X-ray diffraction (XRD) and scanning electron microscopy (SEM). Adhesion strength of the coatings was studied with scratch tests, while antibacterial activity was tested Escherichia coli (E. coli) at various incubation times. Biocompatibility was investigated by osteoblast adhesion. An increase in adhesion strength was observed after heat treatment due to ionic diffusion at the interface. The Ag-incorporated HA coatings killed all E. coli after 24 h of incubation, while no antibacterial activity was observed with pure HA. Ag-incorporated HA coatings had osteoblast adhesion similar to pure HA, which indicated good cytocompatibility. The results of this study show that Ag-incorporated HA is a promise for numerous medical applications. Figure 20 shows osteoblast adhesion results.

Figure 20. Comparison of osteoblast adhesion on glass, pure HA (A0) and Ag-incorporated HA (A2 and A3) bulk samples heat treated at 900°C after 4 h of incubation (the statistical analysis of these data exhibited no significant difference) [47].

Electrospinning

Electrospinning is a synthesis technique driven by voltage and governed by a specific electrohydrodynamic phenomenon in which small fibers are pumped from a polymer solution. This process consists of a solution contained in a reservoir (typically a syringe) tipped with a blunt needle, a pump, a high-voltage power source and a collector. An electric field is established between the needle tip and the collector by applying a voltage. The pump causes the solution to flow at a constant rate, but as charges accumulate at the surface of the liquid, the electrostatic repulsion becomes larger than the surface tension, which results in the liquid meniscus deforming into a conically shaped structure (Taylor cone). The charged liquid then ejects towards the collector, which differs by their geometry (flat plates, rotating drums, mandrels, disks, etc.). If the viscosity of the solution is within a certain range, the technique generates solid fibers as the solvent evaporates [48]. Figure 21 shows a schematic of this technique.

Figure 21. Schematic of the principles of electrospinning [49].

Furko et al. [50] deposited bioactive ions (Mg, Sr, Zn) doped carbonated HA and pure carbonated HA-loaded biopolymer polyvinylpyrrolidone (PVP) coatings using electrospinning and analyzed their morphology. SEM analysis showed that HA particles can be successfully incorporated into the polymeric matrix and simultaneously the bioceramic powders were attached to the surface of polymer fibers. The surface of the fibers was not fully covered with the particles. When the ceramic powder was added to the polymeric solution, the polymer fibers became more entangled, and the diameters of fibers varied over a wide range compared to the polymer fibers without powder addition. Figure 22 shows SEM elemental mapping of the samples.

Figure 22. SEM image on ion-doped Hap-PVP composite with the corresponding elemental maps [50].

Ahmadi et al. [51] obtained triple polyvinylalcochol (PVA) coatings containing HA and 8-hydroxyquinoline (8-HQ) on magnesium substrates by two different methods: electrospinning and immersion. The formation of HA and Mg(8-HQ)₂ was indicated by X-ray diffraction (XRD) and fluorescent images. Electrospinning coating containing HA and 8-HQ showed greater anticorrosion performance compared to the same composition immersion samples. This increment is caused by the release of HA and 8-HQ from the nanofibers. Bacterial infections in magnesium implants have been reduced (reduction of 90% in coatings prepared by electrospinning method) by synthesizing the triple coating containing 8-HQ as an antibacterial compound. Figure 23 shows FTIR spectrums of the samples.

Figure 23. FTIR spectrums of PVA, PVA(HA), PVA(HA)(HQ), 8-HQ [51].

Dejob et al. [52] coated carbonated calcium phosphate apatite (already reported to be osteoconductive and present in bone tissue) shaped in the form of micro-sized filaments via the electrospinning process, in order to mimic the collagen fibrils present in the bone extracellular matrix. Cellular assays with MG-63 show that the chemical and morphological properties of the coated implants promote cell viability and proliferation on them. Figure 24 shows the evaluation of cell viability and proliferation.

Figure 24. Evaluation of cell viability and proliferation by LIVE/DEAD imaging (after one day of culture) and Alamar Blue assay (after one to seven days of culture). MG-63 cells were grown in either direct or indirect contact with the dental implant. For Alamar Blue analysis, all the means (n = 8) are statistically different (p < .0001, test ANOVA followed by Tukey’s post-hoc test) except for the couple shown in the graph (marked as ‘ns’, p = .10) [52].

Spray pyrolysis

Spray pyrolysis technique involves three main stages: precursor solution composition, aerosol generation and transport, and synthesis process (chemical reactions). Every stage is adjusted based on the final characteristics of the target material. At the first stage, the precursor solution will have to contain a compound(s) that will produce the chemical composition required after the pyrolysis stage (formation of the droplets). The selection of the solvent will limit the maximum concentration of the precursor compound and will determine the best choice for the aerosol generation and transport processes, the temperature, and rate of synthesis. At the second stage, the aerosol droplet size distribution, which is determined by the aerosol generation mechanism, will set the morphology of the final material, as well as the proper range of synthesis temperatures. The nature of the carrier gas and the flux rate will increase or reduce the probability of a reaction with the precursor compound. At the last stage, whether the final chemical reaction takes place on a gas phase or on a hot substrate will determine if the final material is a powder or a film coating. The most important synthesis parameters in this technique are the concentration molarity of the precursor solution, the carrier gas flux rate and the synthesis temperature [53]. Figure 25 shows a schematic of this technique.

Figure 25. Scheme of spray pyrolysis technique [54].

Sivaraj et al. [55] deposited Cu-hydroxyapatite/functionalized multiwall carbon nanotube (HA/f-MWCNT) films on 316L stainless steel (SS) using spray pyrolysis and analyzed their crystallinity, surface morphology, elemental composition, corrosion and biological properties. XRD and SEM analysis showed that the coatings exhibited high crystallinity and spherical-shaped morphology. Corrosion analysis revealed that the coatings provided higher barrier properties which is beneficial to achieve higher corrosion protection of 316L SS implant. These coatings also showed better antibacterial ability than coatings non-doped with Cu. The antibacterial ability of these coatings was effective against E. coli compared with other microorganisms. Figure 26 shows the antibacterial ability results.

Figure 26. Zone of inhibition of different organisms treated with hydroxyapatite/f-MWCNT nanocomposite. All the results were expressed in terms of mean and standard deviations. Significance was considered 95% (*p < .05, **p < .01 and ***p < .001) [55].

Ye et al. [56] obtained HA coatings on 316L stainless steel substrates using ultrasonic spray deposition. Continuous and pulsed spray patterns were used coupled with precise control of substrate movement. The ability to form continuous, dense, well-bonded coatings was related primarily to the control of the flowrate, temperature and movement of the substrate. Coatings prepared by continuous spray had relatively large cracks while coatings obtained by pulse spray exhibited finer cracks and higher bonding strength. Figure 27 SEM images of the samples.

Figure 27. Surface morphology of USP coatings prepared with pulsed (a and b) and continuous (c and d) [56].

Aguilar-Frutis et al. [57] coated Ti6Al4V sheet substrates with HA using spray-pyrolysis technique. The chemical composition, surface morphology and structure of the coatings were investigated by XPS, SEM and XRD. XPS confirmed the presence of Ca, P and O as the main elements with a Ca/P ratio of 1.3 XRD analysis showed preferred (113) and (321) crystal orientations in the coatings. Figure 28 shows the X-ray diffraction pattern of sample #10.

Figure 28. X-ray diffraction pattern of sample # 10 [57].

Sol–gel

Sol–gel is a deposition technique which begins with the formation of a ‘sol’, which is a stable dispersion of colloidal particles in a solvent. A ‘gel’ is formed by a three-dimensional continuous network or by the joining of polymer chains. In a colloidal gel, the network is built from agglomerates of colloidal particles, whereas in a polymer gel, the particles have a polymeric substructure composed of aggregates of sub-colloidal particles. Most of the times, van der Waals forces or hydrogen bonds dominate the interactions between the particles. During the synthesis, in most gel systems, covalent interactions dominate, and the gel process is irreversible. The gelation process may be reversible if there are other interactions involved [58]. Figure 29 shows a schematic of this technique.

Figure 29. Stages of the sol–gel process [59].

Ansari et al. [60] deposited polycaprolactone (PCL) and polycaprolactone/fluoride substituted-hydroxyapatite (PCL/FHA) nanocomposite coatings at different compositions (10, 20 and 30 wt.% of FHA) on alkali-treated Ti6Al4V substrates and analyzed their morphology, composition, corrosion and in vitro bioactivity. SEM and XRD analysis showed that sol–gel can create compact and crack-free coatings with homogenous dispersion of FHA in the polymeric matrix. Contact angle analysis showed that when the FHA content increased, the hydrophilicity, adhesion strength and surface roughness increased. By increasing the FHA content and the incubation time in simulated body fluid (SBF), the in vitro bioactivity of the coatings improved. Corrosion analysis showed that the corrosion resistance of the PCL/10wt.% FHA coatings increased compared to pure PCL and PCL/FHA coatings with 20 and 30 wt.% of FHA. Finally, the MTT assay confirmed that the coatings had no cytotoxicity effect and PCL/20wt.% FHA coating promoted the proliferation of MG63 cells. Figure 30 shows the cell viability results.

Figure 30. MTT assay on the pure PCL and PCL/FHA coated specimens consisting of different amounts of FHA (#: P ≤ .05) [60].

Ahmadi et al. [61] obtained HA/TiO₂/Al₂O₃ nanocomposite coatings on 316L SS substrates using dip sol–gel method sintered at 550°C. HA/TiO₂/Al₂O₃ reinforced coatings had the lowest cracks and porosity and had the highest bond strength to the 316L SS substrates (26.5 MPa). The corrosion current density of nanocomposite coatings (HA + 40%wt Al₂O₃ + TiO₂) was 0.091 μA/cm², which is pretty low compared to the values obtained for uncoated substrates (40.35 μA/cm²) and HA coatings (25.26 μA/cm²). The EIS test results showed that in Nyquist curves, the semicircular diameter was more extensive for nanocomposite coatings, which indicates that these coatings had higher corrosion resistance. Cell culture studies showed that the reinforced coatings were non-cytotoxic to human MG63 cell line (HA + 30%wt Al₂O₃ + TiO₂) and that these coatings have a higher ability to form a bone-like layer than the other coatings. However, reinforced coatings with different composition (HA + 40%wt Al₂O₃ + TiO₂) showed lower biocompatibility. Thus, HA + 30%wt Al₂O₃ + TiO₂ coatings have higher biocompatibility, adhesion strength and corrosion resistance than other coatings, so a new turning point for bio-composite coatings has been reached. Figure 31 shows cell viability results.

Figure 31. Viability of MG63 cells [61].

Priyadarshini et al. [62] coated Ti6Al4V with cerium-incorporated HA using sol–gel technique. XRD, SEM and EDS were used to characterize the phase formation and surface morphology of the developed Ce-HA coatings. In vitro study evaluated in SBF indicates good layer formation on the coatings. The cell viability studies using MG-63 cells showed good cell proliferation and attachments on coated implants. EIS was used to study electrochemical properties of coatings and the results showed anti-corrosion property. This study showed that cerium-incorporated HA promotes bioactivity, rapid osseointegration and higher corrosion resistance that makes it a promise for desired biomedical applications. Figure 32 shows cell viability results.

Figure 32. The percentage of cell viability [62].

Pulser-laser deposition

Pulsed laser deposition is a physical vapor deposition technique in which a high-energy laser is focused on a target material located in a vacuum chamber to convert it to a plasma. The deposition can occur under a flow of a certain gas or under ultra-high vacuum conditions. This technique is divided in four stages. In the first stage, the laser beam is absorbed by the target material. A laser with sufficient high energy density and short pulse duration, the laser causes the heat up of all elements in the target, so they are dissociated from the target by collisional, thermal and electronic excitation, and exfoliation and ablated out with the stoichiometry of the target. In the second stage, the ablated materials tend to move towards the substrate and show the forward peaking phenomenon, creating a plasma plume. In the third stage, the energetic particles sputter some of the atoms deposited on the surface and a region where the incident flow and the sputter atoms collide is formed. Film growth happens directly next to this thermalized region, which condensates the energetic particles, so they can be deposited on the substrate. In the fourth stage, the nucleation process is determined by the interfacial energies between the substrate, condensing material and the vapor. Deposition rate and substrate temperature greatly affect the critical size of the nucleus. The crystalline film growth is determined by the surface mobility of the atoms, which diffuse on the surface before getting a stable position, dependent on the temperature of the substrate. High deposition temperatures result in defect-free crystal growth, while low deposition temperatures result in poor crystalline or even amorphous structure [63,64]. Figure 33 shows a schematic of this technique.

Figure 33. Diagram of the pulsed laser deposition chamber [65].

Gonzalez-Estrada et al. [65] deposited HA on Ti6Al4V substrates using pulsed-laser deposition technique. The average hardness and microhardness were analyzed with microindentation and indentation tests. The morphology and chemistry of the coating were analyzed using scanning electron microscopy and X-ray photoelectron spectroscopy. Results showed that high incident energy values result in a more homogeneous morphology with smaller grain size and higher hardness. The chemical composition of the coatings had good stoichiometry. Table 5 shows the chemical composition of the coating from the EDX spectra for the five specimens.

Table 5. Chemical composition of the coatings from the EDX spectra for S1 (sample coated with an incident energy of 248 mJ), S2 (sample coated with an incident energy of 220 mJ), S3 (sample coated with an incident energy of 200 mJ), S4 (sample coated with an incident energy of 180 mJ), S5 (sample coated with an incident energy of 160 mJ) [65].

Element	S1 (wt%)	S2 (wt%)	S3 (wt%)	S4 (wt%)	S5 (wt%)
C	3 ± 1	3 ± 1	3 ± 1	3 ± 1	3 ± 1
O	34 ± 11	40 ± 5	39 ± 1	46 ± 3	28 ± 3
P	21 ± 4	19 ± 1	19 ± 1	18 ± 1	20 ± 1
Ca	42 ± 8	47 ± 4	39 ± 1	33 ± 3	42 ± 1
Al	–	<1	<1	–	3 ± 3
Si	<1	<1	–	–	–
Ca/P	1.9	2.5	2.0	1.9	2.0

Dhinasekaran et al. [66] obtained nanostructured bioactive glass and hydroxyapatite coatings on titanium substrates using pulsed-laser deposition and evaluated their biocompatibility. Raman and IR spectroscopic techniques based on silica and phosphate functional groups mapping confirmed homogeneity in the coatings. Hydroxyapatite and bioactive glass show good hemocompatibility in powder form. Hemocompatibility and cytocompatibility results show that hydroxyapatite coatings are a better option. Figure 34 shows hemocompatibility results of the samples.

Figure 34. Hemocompatibility of bioactive glass and HA has been evaluated as per ASTM F756 at different concentrations, for the Hemolysis %. The graph is marked with symbols as (*) for non-hemolytic samples (0%–2% lysis) and as (+) for slightly hemolytic samples (2%–5% lysis) [66].

Vagolu et al. [67] coated Ti6Al4V with wollastonite and boron nitride-doped HA using pulsed laser deposition. Wear and corrosion resistance of the coatings were analyzed in SBF to simulate human body environment. HA was used to form a bone-like layer, wollastonite to enhance bone-tissue regeneration and boron nitride for its bone-tissue healing and antibacterial properties. The results showed that wear and corrosion resistance of the coatings increased when pulsed laser deposition was used. The best wear resistance was obtained with boron nitride and wollastonite doped HA coatings, while the best corrosion resistance was achieved with coatings of pure HA. Figure 35 shows graphs of friction of the samples versus time for different conditions.

Figure 35. Friction coefficient versus time graphs under (a) dry, (b) SBF and (c) dry + SBF conditions [67].

Drop casting

Drop casting is a deposition technique used to deposit small coatings on small surfaces (∼1 cm²) and requires a very small amount of solvent. A solution in form of drops is poured onto the substrate and is allowed to dry under controlled conditions (pressure and temperature) without spreading. When the droplets are deposited onto the substrate, the liquid first spreads on the surface from the drop locations due to the interfacial forces that drive the droplet outward. As multiple droplets are deposited onto the substrate surface, their edges come in contact with each other, mix and form a noncircular drop. Noncircular drops have uneven deposition rates, highly convex regions, stronger evaporation flux and then denser deposits. The films produced by this method are nonuniform because of the inconsistent drying conditions and uneasy control. These films are thicker at the center and thinner at the edges. Film thickness depends on the volume of dispersion and the particle concentration, both of which can be easily varied. Most of the times, it is desirable to use solvents that are volatile, wet the substrate, and are not susceptible to film instabilities [68,69]. Figure 36 shows a schematic of this technique.

Figure 36. Schematic of the Drop Casting deposition technique [68].

Rios-Pimentel et al. [70] used drop casting to deposit coatings of nanocrystalline HA and amphiphilic peptide nano particles (APNPs) on poly(2-hydroxyethyl methacrylate) (pHEMA) on Ti6Al4V substrates and analyzed their morphology and cell adhesion. SEM images showed that the coatings were homogeneous and all over the substrates. Cell studies showed that the difference between cell adhesion of the coatings made of pHEMA hydrogels with nano HA and APNPs and cell adhesion of the rest of the coatings (pHEMA, PHEMA + HA, pHEMA + APNPs) was statistically significant with a p < .005. Figure 37 shows the cell adhesion results.

Figure 37. Osteoblast density on pHEMA, APNPs and nanocrystalline HA-coated titanium. Data are mean ± SEM; n = 3. p = .005 for all comparisons [70].

Conclusions

Aging of the population, early burn out and lack of exercise are the main reasons why the number of orthopedic surgeries is increasing. Metallic materials are the number one choice to fabricate orthopedic implants due to their mechanical properties, being the most popular stainless steel, titanium alloys and cobalt-chromium. The most common causes of orthopedic implant failure after implantation are infections, corrosion, stress shielding and excessive wear. To address the problem associated with the surface of implant materials, changes related to geometry and coating surface have been developed. Among these methods, coating with HA is an effective method to improve the performance of implant materials. In this article, a comprehensive review of recent studies has been carried out to summarize the impact of HA as a coating material for metallic implants. The antibacterial characteristics of HA, their biocompatibility, corrosion behavior and mechanical properties for performance evaluation are briefly summarized. Many coatings techniques such as RF sputtering, microwave-assisted synthesis, plasma spray, laser-induced single-step coating, electrostatic spray-assisted vapor deposition, electrospinning, spray pyrolysis, sol–gel, pulsed-laser deposition and drop casting have been analyzed (presented, reviewed). A significant enhancement in performance is reported for HA coatings in all the cases reviewed here, so in all of them HA coating resulted in a promising candidate for clinical implant applications. Moreover, some studies showed that the incorporation of impurities into HA, such as Mg and Cu, and the mixing with other materials, such as Aloxite, PVP, PCL and APNPs is effective to achieve the required biocompatible, anti-corrosive and mechanical properties. As already mentioned, all HA coatings reviewed here turned out to be good candidates for clinical implant applications; nevertheless, among the techniques, drop casting seems to be the most promising, for it is cheaper to mount and operate than the rest of the techniques. On the other hand, among the other materials added to the HA coatings, APNPs appear as the best candidate since it increased osteoblast adhesion in a statistically significant way. Finally, nanocrystalline HA coatings have shown greater osteoblast adhesion than regular HA coatings.

Disclosure statement

No potential conflict of interest was reported by the author(s).

Data availability statement

The authors confirm that the data supporting the findings of this study are available within the article and its supplementary materials.

Additional information

Funding

This research was supported by CONACYT (CVU# 335685).