Bioactive Calcium Phosphate Coatings for Bone Implant Applications: A Review

Abstract

This review deals with the design of bioactive calcium phosphate coatings deposited on metallic substrates to produce bone implants. The bioceramic coating properties are used to create a strong bonding between the bone implants and the surrounding bone tissue. They provide a fast response after implantation and increase the lifespan of the implant in the body environment. The first part of the article describes the different compounds belonging to the calcium phosphate family and their main properties for applications in biomaterials science. The calcium-to-phosphorus atomic ratio (Ca/P)_at. and the solubility (K_s) of these compounds define their behavior in a physiological environment. Hydroxyapatite is the gold standard among calcium phosphate materials, but other chemical compositions/stoichiometries have also been studied for their interesting properties. The second part reviews the most common deposition processes to produce bioactive calcium phosphate coatings for bone implant applications. The last part describes key physicochemical properties of calcium phosphate coatings and their impact on the bioactivity and performance of bone implants in a physiological environment.

1. Introduction

The ageing of the world’s population is creating an increasing clinical demand for skeletal repair [1,2,3,4,5,6]. In particular, orthopedic and dental surgeries require metallic bone implants made of titanium alloys [7,8,9,10,11,12,13,14,15,16], steels and iron-based alloys [17,18,19,20,21], or CoCr alloys [22,23,24,25,26,27,28]. The mechanical properties of these alloys are appropriate for load-bearing applications, and they are biocompatible with the body environment. According to the International Union of Pure and Applied Chemistry (IUPAC), biocompatibility is the ability of a material to be in contact with a biological system without producing an adverse effect [29,30,31,32,33,34]. However, surface modification of these metallic bone implants with a coating is necessary to make them bioactive in the body environment. Bioactivity is the property of materials to develop a direct, adherent, and strong bonding with the bone tissue [35,36,37,38,39,40]. Among the bioactive materials, calcium phosphates are most frequently used in industry and academic research. They are ceramic materials with a chemical composition akin to bone mineral, the inorganic component of our bones [41,42,43,44,45,46,47]. Inside the body, their bioactivity confers long-term performance on the metallic bone implant. They prevent bone anchorage failure and delay revision surgery [48,49,50,51,52]. Several methods can be used to produce calcium phosphate coatings on metallic bone implants including plasma spraying, magnetron sputtering, pulsed laser deposition, electrospray deposition, electrophoretic deposition, biomimetic deposition, a sol–gel process combined with dip or spin coating, electrodeposition, and hydrothermal synthesis [53]. Among them, plasma spraying is the main industrial process, extensively used since the 1970s to coat metallic bone implants [54]. Other deposition processes have been developed for decades, and their advantages and drawbacks are well established today. The properties of a calcium phosphate coating depend on the process used to produce it and on the experimental conditions and deposition parameters as well. These are of great importance because the modification of coating properties is known to influence the surface bioactivity of the bone implant in a physiological environment.

2. Calcium Phosphates

Calcium phosphate bioceramics are materials made of calcium ions (${\mathrm{Ca}}^{2+}$) and phosphate ions (${\mathrm{H}}_2{\mathrm{PO}}_4^{-}$, ${\mathrm{HPO}}_4^{2-}$, or ${\mathrm{PO}}_4^{3-}$). Several compounds belong to this family, with different stoichiometries and different phosphate species. They are specifically identified in biomaterials science by their calcium-to-phosphorus atomic ratio (Ca/P)_at. (

Table 1. Calcium phosphates described in the literature as coatings for bone implants.

(Ca/P)at.	Calcium Phosphate	Abbreviation	Chemical Formulae	Solubility [-log Ks]	References
2.00	tetracalcium phosphate	TTCP	${\mathrm{Ca}}_4{({\mathrm{PO}}_4)}_2{\mathrm{O}}_2$	38.0–44.0	[55,56,57]
1.67	hydroxyapatite	HAP	${\mathrm{Ca}}_{10}{({\mathrm{PO}}_4)}_6{\mathrm{OH}}_2$	116.8	[58,59,60]
1.50	α-tricalcium phosphate	α-TCP	$\mathsf{\alpha }-{\mathrm{Ca}}_3{({\mathrm{PO}}_4)}_2$	25.5	[61,62,63]
1.50	β-tricalcium phosphate	β-TCP	$\mathsf{\beta }-{\mathrm{Ca}}_3{({\mathrm{PO}}_4)}_2$	28.9	[64,65,66]
1.34–1.66	calcium-deficient apatite	Ca-def apatite	${\mathrm{Ca}}_{10-\mathrm{x}}{({\mathrm{HPO}}_4)}_{\mathrm{x}}{({\mathrm{PO}}_4)}_{6-\mathrm{x}}{(\mathrm{OH})}_{2-\mathrm{x}}$ with 0 < x < 2	85.1	[67,68,69]
1.33	octacalcium phosphate	OCP	${\mathrm{Ca}}_8{({\mathrm{HPO}}_4)}_2{({\mathrm{PO}}_4)}_4·5{\mathrm{H}}_2\mathrm{O}$	96.6	[70,71,72]
1.00	calcium pyrophosphate	CPP	${\mathrm{Ca}}_2{\mathrm{P}}_2{\mathrm{O}}_7$	18.5	[73,74,75]
1.00	dicalcium phosphate anhydrous, also known as monetite	DCPA	${\mathrm{CaHPO}}_4$	6.9	[76,77,78]
1.00	dicalcium phosphate dihydrate, also known as brushite	DCPD	${\mathrm{CaHPO}}_4·2{\mathrm{H}}_2\mathrm{O}$	6.6	[79,80,81]
0.50	monocalcium phosphate anhydrous	MCPA	$\mathrm{Ca}{({\mathrm{H}}_2{\mathrm{PO}}_4)}_2$	1.1	[82,83,84]
0.50	monocalcium phosphate monohydrate	MCPM	$\mathrm{Ca}{({\mathrm{H}}_2{\mathrm{PO}}_4)}_2·{\mathrm{H}}_2\mathrm{O}$	1.1	[85,86,87]

).

The stoichiometry of a calcium phosphate coating affects its solubility in a physiological environment, which is the first step involved in the bioactivity process after implantation (Figure 1).

The partial dissolution of the calcium phosphate coating in contact with the physiological environment induces ionic releases. The local concentrations of calcium and phosphate ions increase up to supersaturation, which triggers the precipitation of biological apatite at the interface between the implant and the surrounding bone tissues [30,31,35,36,37,38]. After these first chemical steps, the biological steps start, involving bone cell attachment, proliferation, and differentiation. In the last step of the bioactivity process, the bone cells trigger the formation of the extracellular matrix (ECM), which is a three-dimensional network of macromolecules and minerals, such as collagen, enzymes, glycoproteins, and apatite [88,89,90]. The function of the extracellular matrix is to provide structural and biochemical support to the surrounding bone cells to promote their development [91]. Due to the bioactivity of the calcium phosphate coatings, bone-like apatite is formed at the interface between the bone implant and the bone tissue. This bone-like apatite layer is a direct, adherent, and strong bonding that results in the long-term stability of the bone implant inside the human body [92]. However, the success of the bioactivity process is related to several properties of the calcium phosphate coating and not only to the stoichiometry and solubility of the bioceramic material. The choice of the process and the experimental deposition conditions may influence many physicochemical properties of the calcium phosphate coating, and consequently the bioactivity process.

3. Deposition Methods

3.1. Plasma Spraying (PS)

Plasma spraying is the most widespread industrial process because it is remarkably efficient at producing large quantities of bioceramic coatings on metallic bone implants. However, a perfect reproducibility of the properties of the deposited coatings is impossible to achieve because of the highly nonlinear nature of the process [93,94].

In atmospheric plasma spraying (APS), calcium phosphate powder (generally hydroxyapatite) is injected into a plasma jet, the temperature of which is in the range of ten thousand degrees [95,96]. At this high temperature, the grains of powder are molten or partly molten. The plasma jet directs the molten droplets toward the bone implant surface, where the steps of spreading, accumulation, cooling, and solidification produce a coating (Figure 2).

However, the high temperatures of the process give rise to several issues. The calcium phosphate particles melt incongruently, locally resulting in structural modifications, uncontrolled phase changes, and chemical decompositions. These modifications produce a coating, the physicochemical and biological properties of which differ from those of the initial powder [98,99,100]. The thermal decomposition of hydroxyapatite within a plasma is comprehensively described by Heimann’s work on the following reactions [101]:

$${\mathrm{Ca}}_{10}{\left({\mathrm{PO}}_4\right)}_6{\left(\mathrm{OH}\right)}_2\to {\mathrm{Ca}}_{10}{\left({\mathrm{PO}}_4\right)}_6{\left(\mathrm{OH}\right)}_{2-2\mathrm{x}}{\mathrm{O}}_{\mathrm{x}}{⬚}_{\mathrm{x}}+{\mathrm{xH}}_2\mathrm{O}$$

(1)

$${\mathrm{Ca}}_{10}{\left({\mathrm{PO}}_4\right)}_6{\left(\mathrm{OH}\right)}_{2-\mathrm{x}}{\mathrm{O}}_{\mathrm{x}}{⬚}_{\mathrm{x}}\to {\mathrm{Ca}}_{10}{\left({\mathrm{PO}}_4\right)}_6{\mathrm{O}}_{\mathrm{x}}{⬚}_{\mathrm{x}}+\left(1-\mathrm{x}\right){\mathrm{H}}_2\mathrm{O}$$

(2)

$${\mathrm{Ca}}_{10}{\left({\mathrm{PO}}_4\right)}_6{\mathrm{O}}_{\mathrm{x}}{⬚}_{\mathrm{x}}\to 2{\mathrm{Ca}}_3{\left({\mathrm{PO}}_4\right)}_2+{\mathrm{Ca}}_4\mathrm{O}{\left({\mathrm{PO}}_4\right)}_2$$

(3)

$${\mathrm{Ca}}_3{\left({\mathrm{PO}}_4\right)}_2\to 3\mathrm{CaO}+{\mathrm{P}}_2{\mathrm{O}}_5$$

(4)

$${\mathrm{Ca}}_4\mathrm{O}{\left({\mathrm{PO}}_4\right)}_2\to 4\mathrm{CaO}+{\mathrm{P}}_2{\mathrm{O}}_5$$

(5)

As a function of the experimental parameters, these five reactions may occur during plasma spray deposition, where ${⬚}_{\mathrm{x}}$ refers to lattice vacancies in the crystal structure of the calcium phosphate compound. The resulting bioceramic coating contains a mixture of oxyhydroxyapatite (${\mathrm{Ca}}_{10}{\left({\mathrm{PO}}_4\right)}_6{\left(\mathrm{OH}\right)}_{2-2\mathrm{x}}{\mathrm{O}}_{\mathrm{x}}{⬚}_{\mathrm{x}}$), oxyapatite (${\mathrm{Ca}}_{10}{\left({\mathrm{PO}}_4\right)}_6{\mathrm{O}}_{\mathrm{x}}{⬚}_{\mathrm{x}}$), tricalcium phosphate (${\mathrm{Ca}}_3{\left({\mathrm{PO}}_4\right)}_2$), tetracalcium phosphate (${\mathrm{Ca}}_4\mathrm{O}{\left({\mathrm{PO}}_4\right)}_2$), and calcium oxide ($\mathrm{CaO}$) instead of pure hydroxyapatite as initially expected. All these additional phases affect the physicochemical properties of the coatings. Moreover, atmospheric plasma spray deposition produces coatings with residual stress, cracks, and interconnected porosity. These are caused by differences in the coefficients of thermal expansion of the substrate and coating, the imperfect melting of the particles, the insufficient flow of molten droplets in contact with the substrate, a rapid solidification rate, and poor interlayer bonding [102,103]. The rapid solidification induces a local melt-quenching of the particles that results in the amorphization of the bioceramics. The control of the chemical composition and structural properties of plasma-sprayed calcium phosphate coatings is difficult. They are made of several phases in several crystalline states, resulting in a highly heterogeneous bioactive behavior in a physiological environment. Nonetheless, the process is efficient in reaching industrial objectives, i.e., the production of large quantities of coatings at a low cost. The mechanical properties of the coatings are also satisfactory, especially their hardness and long-term stability in normal storage conditions. The adhesion to the metallic substrate is generally high enough, even though many research studies are still trying to find solutions to improve it [104]. Adhesion is a key property of industrial calcium phosphate coatings, the value of which is standardized for the biomedical market (see Section 4.6).

Atmospheric plasma spraying needs good flowability of the injected powder. This requirement limits the particle size of hydroxyapatite to coarse grains in the range of tens to hundreds of micrometers. Submicrometric powders cannot be directly used, because they tend to agglomerate readily due to high surface energy. The resulting flowability is not appropriate for plasma spray deposition. Suspension plasma spraying (SPS) and solution precursor plasma spraying (SPPS) are recent alternative processes that use a liquid feedstock injected into the plasma jet to produce sprayed calcium phosphate coatings [100]. Water or ethanol, or a mixture of both, is generally used. These two processes can produce nanostructured calcium phosphate coatings noted for their enhanced osseoconductive behavior [105,106].

Because the plasma spraying processes have advantages and drawbacks, the study of alternative deposition methods to produce calcium phosphate coatings for bone implant applications remains a major research topic for academic and industrial biomedical research.

3.2. Magnetron Sputerring (MS)

Magnetron sputtering of a calcium phosphate target is an alternative solution for producing bioactive calcium phosphate coatings on bone implants. Magnetron sputtering is a physical vapor deposition (PVD) process. A deposition chamber at room temperature is evacuated to a high vacuum to remove all potential contaminants. After the base pressure has been reached, a working gas is injected into the chamber, usually a noble gas such as argon. The resulting pressure is typically around 1 Pa. Plasma is then ignited from this noble gas by applying a high voltage between the cathode connected to the target and the anode connected to the deposition chamber as an electrical ground (Figure 3). The voltage necessary to start a discharge in a gas between two electrodes as a function of pressure and gap length is given by Paschen’s law [107,108]. The process requires plasma ignition and a self-sustained discharge. Plasma contains high-energy ions that collide with the atoms of the target with enough energy to eject and transport them toward the surface of the bone implant to progressively form a coating [109].

Direct current (DC) magnetron sputtering cannot be used to sputter insulating materials such as ceramics because of the charge accumulation within the target during the process. Pulsed-DC and radio frequency (RF) magnetron sputtering are alternative solutions for depositing insulating materials [110,111,112]. They produce dense, uniform, and adherent calcium phosphate coatings. However, the different elements of a multicomponent target have different sputtering behaviors. The elemental stoichiometry of the deposited coating usually differs from that of the target. The experimental parameters of the process can be used to modify some properties of the deposited calcium phosphate coatings such as stoichiometry, morphology, and structure, resulting in different bioactive behaviors [113,114,115].

3.3. Pulsed Laser Deposition (PLD)

Pulsed laser deposition is another PVD process carried out in a vacuum chamber [116,117,118]. The ablation of a calcium phosphate target hit by a high-power laser produces a plasma plume composed of ejected atoms, ions, and electrons (Figure 4). In contact with the substrate, the ejected material nucleates and grows to form a surface coating.

The efficiency of the process mainly depends on laser beam properties such as wavelength, energy density, fluence, and pulse width. Pulsed laser deposition produces uniform and adherent thin coatings. However, as observed for magnetron sputtering, the elemental stoichiometry of the target and that of the deposited coating are not identical. The physicochemical and biological properties of the coating are impacted by the experimental conditions of the process [120,121,122].

3.4. Electrospray Deposition (ESD)

Electrospray deposition requires a precursor solution containing calcium and phosphate ions, or a suspension of calcium phosphate particles. The solution is sprayed by using a syringe through a nozzle that is connected to a high voltage (Figure 5).

At the end of the capillary tube, the meniscus of the conducting solution becomes conical when charged (Taylor cone). Charged droplets are formed by the continuous breakup of the steady jet of solution leaving the tip of the nozzle. Solvent evaporation on the way toward the bone implant surface promotes the shrinkage of the charged droplets. In contact with the grounded and heated substrate, these very small droplets lose their surface charge and dry, progressively producing the bioactive coating (Figure 6).

The morphology and structure of the coatings are impacted by properties of the solution such as conductivity and surface tension, and by electrospraying parameters such as voltage, flowrate, and distance between the needle tip and the substrate [124,125,126,127,128].

3.5. Electrophoretic Deposition (EPD)

Electrophoretic deposition occurs by means of the migration of calcium phosphate particles in a colloidal suspension [129,130,131]. In a solution, typically water or ethanol, the calcium phosphate particles carry a positive or negative surface charge due to electrostatic interactions with the ionic species of the solution. This surface charge induces the formation of a diffuse double layer containing anions and cations (Figure 7).

The potential difference between the solution and the interface of the two layers is called zeta potential (ζ). This surface potential impacts the stability of colloidal dispersions by inducing electrostatic interactions between the particles of the suspension [132,133,134,135,136,137,138]. Thanks to the zeta potential, the particles can be accelerated under the influence of an electric field between two conductive electrodes connected to a generator. If the particles are positively charged, they move through the liquid toward the cathode (cathodic EPD in Figure 8a). If the particles are negatively charged, they move toward the anode (anodic EPD in Figure 8b).

When a particle reaches the surface of an electrode, the size of the double layer is reduced (Figure 9a), promoting the progressive accumulation and coagulation of particles to form a calcium phosphate coating (Figure 9b).

The main parameters for the success of the EPD process are the pH and the stability of the suspension, the dielectric constant (ε) and the viscosity (η) of the solvent, the average particle size, the substrate conductivity, the voltage, the distance between the electrodes, and the deposition time. Post-deposition thermal annealing is required to evaporate the solvent and to improve the cohesive and adhesive properties of the coating [139,140].

3.6. Biomimetic Deposition

The surface of titanium and titanium alloys is naturally covered by a native oxide layer of TiO₂ produced by their reaction with oxygen in the air [141]. This surface layer is bioactive and promotes the slow deposition of a calcium phosphate coating during immersion in simulated body fluid (SBF) at 37 °C, an acellular solution with pH and ion concentrations similar to those of human blood plasma (

Table 2. Ion concentrations of blood plasma and SBF.

Ion	Concentrations (mM)
	Blood Plasma (7.2 < pH < 7.4)	SBF (pH = 7.4)
${\mathrm{Na}}^{+}$	142.0	142.0
${\mathrm{K}}^{+}$	5.0	5.0
${\mathrm{Mg}}^{2+}$	1.5	1.5
${\mathrm{Ca}}^{2+}$	2.5	2.5
${\mathrm{Cl}}^{-}$	103.0	147.8
${\mathrm{HCO}}_3^{-}$	27.0	4.2
${\mathrm{HPO}}_4^{2-}$	1.0	1.0
${\mathrm{SO}}_4^{2-}$	0.5	0.5

).

These physiological conditions trigger the spontaneous nucleation and growth of apatite on the surface of the native ${\mathrm{TiO}}_2$ layer. In the simulated body fluid, ${\mathrm{OH}}^{-}$ groups of the solution are adsorbed at the surface of ${\mathrm{TiO}}_2$ and bond to titanium ions to produce $\mathrm{Ti}-\mathrm{OH}$ groups. In slightly basic conditions, their deprotonation produces $\mathrm{Ti}-{\mathrm{O}}^{-}$ groups that attract ${\mathrm{Ca}}^{2+}$ ions from the solution to form an amorphous surface layer of calcium titanate. Then, this positively charged layer attracts negatively charged phosphate ions to form a stable phase of amorphous calcium phosphate [142].

The biomimetic deposition process is very slow. Several days or weeks of immersion in SBF are necessary to produce a calcium phosphate coating a few micrometers thick. However, faster depositions have been observed for pretreated titanium surfaces [143]. Thermal annealing in air produces a thicker ${\mathrm{TiO}}_2$ layer, the porosity of which accelerates the biomimetic deposition process. Another relevant pretreatment process is the immersion of titanium in a highly concentrated $\mathrm{NaOH}$ solution (typically 10 M) at 60 °C for 24 h. This alkaline treatment results in the formation of sodium titanate on the surface, which increases the reaction kinetics of the biomimetic deposition (Figure 10).

3.7. Sol–Gel Process Combined with Dip or Spin Coating

Sol–gel is a low-temperature process that transforms an inorganic colloidal suspension (sol) into a three-dimensional network structure containing a liquid phase (gel). Calcium phosphate materials are produced by using calcium and phosphorus precursors [145,146,147]. Examples of calcium precursors described in the literature are calcium acetate monohydrate ($\mathrm{Ca}{\left({\mathrm{CH}}_3\mathrm{COO}\right)}_2·{\mathrm{H}}_2\mathrm{O}$), calcium nitrate tetrahydrate ($\mathrm{Ca}{\left({\mathrm{NO}}_3\right)}_2·4{ \mathrm{H}}_2\mathrm{O}$), and calcium chloride (${\mathrm{CaCl}}_2$). Examples of phosphorus precursors are phosphoric acid (${\mathrm{H}}_3{\mathrm{PO}}_4$), triethyl phosphite ($\mathrm{P}{\left({\mathrm{OCH}}_2{\mathrm{CH}}_3\right)}_3$), phosphorus pentoxide (${\mathrm{P}}_2{\mathrm{O}}_5$), diammonium hydrogen orthophosphate $\left({\left({\mathrm{NH}}_4\right)}_2{\mathrm{HPO}}_4\right)$, and trisodium phosphate (${\mathrm{Na}}_3{\mathrm{PO}}_4$). They are dissolved separately in solvents that are typically water, ethanol, or a mixture of both. The dissolved precursors are then mixed dropwise and gently stirred. The two reactions involved in the process are hydrolysis and condensation, the kinetics of which can be controlled by adjusting the pH value of the solution [148,149].

Thanks to the viscosity of the sol, the sol–gel process can be combined with dip coating or spin coating techniques to produce a surface coating (Figure 11). The dip coating technique involves the immersion and withdrawal of a substrate in the sol. The evaporation of the solvents in the air atmosphere during the drying step triggers the gelation process (polycondensation), which results in the formation of a calcium phosphate coating. The thickness of the coating depends mainly on the withdrawal speed and the viscosity of the sol.

The spin coating technique requires the deposition of the sol onto a substrate that is rotating around an axis perpendicular to the coated surface (Figure 11). The gelation step is also triggered by the evaporation of the solvent. The thickness of the coating depends mainly on the rotational speed and the viscosity of the sol.

The depositions are typically followed by thermal annealing at hundreds of degrees to densify the calcium phosphate coating and improve its mechanical properties.

3.8. Electrochemical Deposition (ECD)

The electrodeposition of calcium phosphate coatings requires two electrodes immersed in an electrolytic solution of calcium and phosphate ions. They are connected to a generator (Figure 12) [150,151,152,153,154,155].

Electrochemical reactions occur at both electrode–electrolyte interfaces. The reduction of water, the solvent of the solution, takes place at the cathode surface as follows:

$$2{\mathrm{H}}_2\mathrm{O}+2{\mathrm{e}}^{-}\to {\mathrm{H}}_2\uparrow +2{\mathrm{OH}}^{-}$$

(6)

If the solution is acidic, the reduction of protons may also occur at the cathode surface:

$$2{ \mathrm{H}}^{+}+2{\mathrm{e}}^{-}\to {\mathrm{H}}_2\uparrow$$

(7)

The resulting local pH variation triggers the precipitation of a calcium phosphate coating (Figure 13) [156,157,158,159,160,161].

The chemical composition and the stoichiometry of the precipitated coating depend on the pH value at the cathode, which is impacted by the process parameters. The following phases can be obtained:

-

dicalcium phosphate dihydrate (brushite):

Ca²⁺ + HPO₄²⁻ + 2H₂O → CaHPO₄∙2H₂O

(8)

-

octacalcium phosphate:

$$8{\mathrm{Ca}}^{2+}+2{\mathrm{HPO}}_4^{2-}+4{\mathrm{PO}}_4^{3-}+5{\mathrm{H}}_2\mathrm{O} \to {\mathrm{Ca}}_8{\left({\mathrm{HPO}}_4\right)}_2{({\mathrm{PO}}_4)}_4·5{\mathrm{H}}_2\mathrm{O}$$

(9)

-

calcium-deficient apatite:

$$\left(10-\mathrm{x}\right){\mathrm{Ca}}^{2+}+{\mathrm{xHPO}}_4^{2-}+\left(6-\mathrm{x}\right){\mathrm{PO}}_4^{3-}+\left(2-\mathrm{x}\right){\mathrm{OH}}^{-} \to {\mathrm{Ca}}_{10-\mathrm{x}}{\left({\mathrm{HPO}}_4\right)}_{\mathrm{x}}{({\mathrm{PO}}_4)}_{6-\mathrm{x}}{\left(\mathrm{OH}\right)}_{2-\mathrm{x}}$$

(10)

with 0 < x < 2

-

hydroxyapatite:

$$10{\mathrm{Ca}}^{2+}+6{\mathrm{PO}}_4^{3-}+2{\mathrm{OH}}^{-} \to {\mathrm{Ca}}_{10}{({\mathrm{PO}}_4)}_6{\left(\mathrm{OH}\right)}_2$$

(11)

The first experiments typically used direct current, but pulsed current electrodeposition has become more usual in the most recent years. The break times are used to remove the H₂ bubbles and to homogenize the electrolyte concentrations [162,163,164,165,166].

Another solution for reducing the amount of ${\mathrm{H}}_2$ bubbles is the addition of hydrogen peroxide (${\mathrm{H}}_2{\mathrm{O}}_2$) to the electrolyte solution [153]. Hydrogen peroxide is a strong oxidative reagent, the reduction of which produces hydroxide ions at the cathode according to the reaction (12):

$${\mathrm{H}}_2{\mathrm{O}}_2+2{\mathrm{e}}^{-}\to 2{\mathrm{OH}}^{-}$$

(12)

However, the concentration of hydrogen peroxide in the electrolytic solution is limited because the overproduction of hydroxide ions prevents the deposition of the coating [153]. The optimization of the process showed that pulsed current electrodeposition from a 9 vol% ${\mathrm{H}}_2{\mathrm{O}}_2$ electrolyte solution produces stoichiometric hydroxyapatite (Ca/P = 1.67) according to the reaction (11).

In addition, the ionic substitution of the electrodeposited calcium phosphate coating can be easily obtained by modifying the electrolyte composition. Due to the low temperature of the process, the addition of organic components (polymers, proteins, drugs, etc.) is also possible to improve the biological and mechanical performances of the electrodeposited coating [167,168,169].

3.9. Hydrothermal Synthesis

Hydrothermal synthesis produces crystallized calcium phosphate coatings in a high-temperature solution under high pressure [170,171]. The aqueous solution contains calcium and phosphate ions. The process is carried out in an autoclave (Figure 14), typically at temperatures ranging from 100 °C to 350 °C and pressures up to 10⁷ Pa (100 bar). These extreme experimental conditions induce the precipitation of crystalline calcium phosphate crystals that nucleate and grow on the surface of bone implants. The crystallinity and the morphology of the coating are highly influenced by the pH of the solution and the temperature used during the process. As a function of the experimental conditions, different morphologies can be achieved, such as nanorods, microspheres, flakes, needles, hexagonal prisms, and hollow flowerlike structures [172].

The process can also be used as a complementary step to modify the crystallinity, morphology, porosity, and phase composition of calcium phosphate coatings deposited by other deposition methods, such as electrodeposition or plasma spraying. In fact, most of the previously described deposition methods produce calcium phosphates of low crystallinity. Dense calcium phosphate coatings with improved properties are described in the literature after post-deposition hydrothermal treatment under high temperature and high pressure [173,174]. Several works show the conversion of amorphous DCPD, OCP, or TCP into crystallized stoichiometric HAP by means of hydrothermal treatment.

4. Main Properties Impacting the Bioactivity of Calcium Phosphate Coatings

In addition to stoichiometry and solubility, several physicochemical properties impact the bioactivity of calcium phosphate coatings immersed in a physiological environment. Crystallinity, morphology, roughness, porosity, wettability, adhesion, and ionic substitution are the most important ones.

4.1. Crystallinity

The crystallinity of calcium phosphate coatings impacts their solubility in a physiological environment. The more crystallized the coating, the more stable it is in solution [175,176,177]. Crystallinity can be controlled by post-deposition thermal annealing. The international standard ISO 13779-2 recommends a degree of crystallinity higher than 45% for the biomedical market of bone implants [178]. However, as a function of the annealing temperature, several phases can form in addition to the calcium phosphate phases [179,180]. To maintain a low level of cytotoxicity, the quantity of secondary phases (for example CaO) in the calcium phosphate coatings should be below 5 wt.% [178]. The methods for determining the crystallinity of calcium phosphate coatings and the quantity of secondary phases are comprehensively described in the international standard ISO 13779-3 [181].

4.2. Morphology

The surface morphology of calcium phosphate coatings affects the bone cells’ attachment, growth, proliferation, and differentiation [182,183]. As a function of the deposition process and the experimental conditions, the surface morphology of the coatings can change [184,185]. Regular surface morphologies are more efficient for bone cell attachment than irregular and sharp ones [186]. According to Cairns et al., they significantly promote the expression of growth factors involved in bone formation in comparison with sharp surfaces made of needles [187,188].

4.3. Roughness

Bioactivity is a surface phenomenon influenced by, among other factors, the roughness of materials. High roughness exceeding 2 µm is not appropriate, because the long distances between valleys and peaks prevent the formation of the osteoblastic pseudopodia required for bone cell adhesion [189,190,191]. Calcium phosphate coatings with roughness values in the range of 0.5 to 1.5 μm are generally described to be the most interesting for the promotion of bone cell activity [192,193,194].

4.4. Porosity

The porosity of calcium phosphate coatings has a significant impact on the bioactive behavior of bone implants in a physiological environment. Pores larger than one hundred micrometers (macroporosity) support the growth of bone tissues through the coating and improve the connection between newly formed bone cells. However, these large pores also strongly reduce the mechanical properties of the bioceramic coatings [195]. Smaller pores of a few tens of micrometers and below (microporosity) enhance protein adsorption, body fluid circulation, and the resorption rate of the coating [196].

4.5. Wettability

Surface wettability is a key property of calcium phosphate coatings because the bioactivity processes occur in a liquid medium. Contact angle (θ) measurements are used to quantify the wetting behavior of a drop of physiological solution deposited on the coating surface [197,198,199]. As a function of the contact angle value, the surface is qualified as hydrophilic or hydrophobic (Figure 15).

Biomaterials with hydrophilic surfaces are more effective in promoting chemical and biological interactions with the physiological environment [201,202].

4.6. Adhesion

The adhesion of calcium phosphate coatings is the main mechanical property required by the biomedical market [104,203,204,205,206,207]. The value of coatings is determined by performing tensile adhesion measurements according to the international standard ISO 13779-4 [208].

The measurement requires a Ti6Al4V cylinder (25 mm in diameter and 25 mm in height) with one surface coated with calcium phosphate. The coated surface is attached to another Ti6Al4V cylinder by adhesive glue (Figure 16a). The entire system is introduced into a standard tensile machine where an increasing load is applied (Figure 16b) until the separation of the coating is achieved by the breaking of the interface with the initially coated cylinder (Figure 16c). A cohesive failure inside the coating may also occur, but in this case, the measurement is not valid and must be repeated. A minimum of five measurements of adhesive failure is necessary to obtain an average adhesion value. The bone implant industry requires adhesion values higher than 15 MPa [208].

This protocol is standardized for industrial applications, but several other methods can be used to determine the adhesion of calcium phosphate coatings, including the peel test, the scratch test, the ultrasonic test, and the laser shock adhesion test [104,209,210,211,212,213].

4.7. Ionic Substitution for Biological Enhancement

The bioactivity and biological properties of calcium phosphate coatings can be improved by means of ionic substitution [214,215,216,217,218,219,220,221,222,223]. The objective is to release the substituting ions in the physiological environment after implantation, taking advantage of the dissolution process (see Section 2). Several ionic substitutions have been described in the literature, using monovalent cations, divalent cations, trivalent cations, or anions. They are used to impart the various biological or chemical effects described in

Table 3. Ions used as substituents in calcium phosphate coatings.

Ions	Biological/Chemical Effect	References
monovalent cations
${\mathrm{Ag}}^{+}$	antibacterial activity	[224,225,226]
${\mathrm{K}}^{+}$	osteogenesis	[227,228,229]
${\mathrm{Li}}^{+}$	osteogenesis	[230,231,232]
${\mathrm{Na}}^{+}$	osteogenesis	[233,234,235]
divalent cations
${\mathrm{Co}}^{2+}$	angiogenesis	[236,237,238]
${\mathrm{Cu}}^{2+}$	antibacterial activity	[239,240,241]
${\mathrm{Mg}}^{2+}$	osteogenesis	[242,243,244]
${\mathrm{Mn}}^{2+}$	osteogenesis	[245,246,247]
${\mathrm{Sr}}^{2+}$	osteogenesis	[248,249,250,251]
${\mathrm{Zn}}^{2+}$	osteogenesis/antibacterial/anti-inflammatory	[252,253,254]
trivalent cations
${\mathrm{Bi}}^{3+}$	anticancer/antibacterial	[255,256,257]
${\mathrm{Ce}}^{3+}$	antibacterial	[258,259,260,261]
${\mathrm{Er}}^{3+}$	photoluminescence	[262,263,264]
${\mathrm{Eu}}^{3+}$	osteogenesis/photoluminescence	[265,266,267]
${\mathrm{Fe}}^{3+}$	osteogenesis/anticancer/antibacterial	[268,269,270]
${\mathrm{Ga}}^{3+}$	anticancer/antibacterial	[271,272,273]
${\mathrm{Tb}}^{3+}$	photoluminescence	[274,275]
anions
${\mathrm{Cl}}^{-}$	osteogenesis	[276,277,278]
${\mathrm{CO}}_3^{2-}$	osteogenesis	[279,280,281]
${\mathrm{F}}^{-}$	antibacterial	[282,283,284]
${\mathrm{SeO}}_3^{2-}$/${\mathrm{SeO}}_4^{2-}$	anticancer/antibacterial	[285,286,287]
${\mathrm{SiO}}_4^{4-}$	osteogenesis	[288,289,290]

.

A few percent of these ions are generally used to produce substituted calcium phosphate coatings. Multi-substitution with several substituting ions is also described in the literature with the objective of cumulating its positive effects on the biological properties of bone implants [291,292,293,294,295,296,297,298,299,300,301,302,303,304,305].

5. Conclusions

This article reviewed the calcium phosphate compounds that are used as coatings to make the surface of metallic bone implants osseoconductive. The link between the stoichiometry, solubility, and bioactivity of calcium phosphate coatings was explained. The main processes used in industry and academic research to design calcium phosphate coatings were described. Historically, plasma spraying was the first industrial process, but interesting alternative methods were also developed and have been described herein. The stoichiometry and the physicochemical properties of the calcium phosphate coatings depend crucially on the deposition process and the experimental parameters used during coating deposition. The impact of coating properties on bioactivity has been briefly described. Finally, the ionic substitution of calcium phosphate coatings was reviewed from the literature, including the biological enhancements provided by ionic substitution.