Electrophoretic Deposition of Bioactive Glass Coatings for Bone Implant Applications: A Review

Abstract

This literature review deals with the electrophoretic deposition of bioactive glass coatings on metallic substrates to produce bone implants. Biocompatible metallic materials, such as titanium alloys or stainless steels, are commonly used to replace hard tissue functions because their mechanical properties are appropriate for load-bearing applications. However, metallic materials barely react in the body. They need a bioactive surface coating to trigger beneficial biological and chemical reactions in the physiological environment. Bioactive coatings aim to improve bone bonding, shorten the healing process after implantation, and extend the lifespan of the implant. Bioactive glasses, such as 45S5, 58S, S53P4, 13-93, or 70S30C, are amorphous materials made of a mixture of oxides that are accepted by the human body. They are used as coatings to improve the surface reactivity of metallic bone implants. Their high bioactivity in the physiological environment induces the formation of strong chemical bonding at the interface between the metallic implant and the surrounding bone tissue. Electrophoretic deposition is one of the most effective solutions to deposit uniform bioactive glass coatings at low temperatures. This article begins with a review of the different compositions of bioactive glasses described in the scientific literature for their ability to support hard tissue repair. The second part details the different stages of the bioactivity process occurring at the surface of bioactive glasses immersed in a physiological environment. Then, the mechanisms involved in the electrophoretic deposition of bioactive glass coatings on metallic bone implants are described. The last part of the article details the current developments in the process of improving the properties of bioactive glass coatings by adding biocompatible elements to the glassy structure.

1. Introduction

The clinical demand for skeletal repair is constantly increasing due to the aging of the world population [1,2,3,4,5,6,7,8]. Metallic bone implants are commonly used to replace deficient hard tissues of the body because their mechanical properties are appropriate for load-bearing applications [9]. The main metallic materials used to produce bone implants are titanium alloys [10,11,12,13], iron-based alloys and stainless steels [14,15,16,17,18,19,20,21], tantalum [22,23,24,25,26,27,28,29,30], and cobalt chromium alloys [31,32,33,34,35,36,37,38]. They are biocompatible with the human body, meaning they do not cause any adverse effects when they are implanted [39,40,41,42,43,44,45]. The mechanical properties of these metallic materials make them relevant in replacing hard tissues, but they barely react with the physiological environment. They need a coating to make their surface bioactive, i.e., able to trigger biological activity and chemical reactions, promoting the formation of an adherent and strong bond with bone tissue (Figure 1). Bioactive coatings deposited on the surface of metallic implants promote long-term stability in the body and prevent or delay revision surgery [46,47,48,49,50,51,52].

Bioactive glasses are the most reactive materials in a physiological environment. The idea to use them for hard tissue repair was introduced first by Larry Hench in 1969, who produced the famous 45S5 Bioglass^® by mixing the four oxides SiO₂, CaO, Na₂O, and P₂O₅ [53,54,55]. In the following years, the ability of 45S5 Bioglass^® to bond to bone tissues was demonstrated in vivo with several animal experiments [56,57,58,59]. For more than five decades, the bioactivities of glasses of other compositions have been deeply studied, and the idea of using them as coatings deposited on metallic bone implants has grown rapidly. Bioactive glass coatings are commonly described as an alternative to calcium phosphate coatings [60,61,62,63,64,65,66,67,68,69,70] because of their high reactivity in physiological environments.

Several methods have been explored to produce bioactive glass coatings on bone implants, such as plasma spraying [71,72,73,74,75,76], laser cladding [77,78,79,80,81,82,83], magnetron sputtering [84,85,86,87,88,89,90,91,92,93,94], sol–gel process combined with dip or spin coating [95,96,97,98,99,100,101,102,103], or electrophoretic deposition [104,105,106,107,108,109,110]. Among these deposition methods, electrophoretic deposition is attracting increasing attention from industrial and academic researchers because the process produces uniform coatings at low temperatures using quite inexpensive and easy-to-implement equipment. The process requires a stable colloidal suspension of bioactive glass particles, the surface of which is electrically charged when it is in contact with the solution. Under the influence of an electric field, the particles move through the liquid toward an oppositely charged electrode, where they progressively accumulate to form a coating [111,112,113]. To better understand the mechanisms involved in the process, this article reviews the scientific literature dealing with the electrophoretic deposition of bioactive glass coatings for bone implant applications. The main bioactive glass materials typically used for bone repair are presented first (

Table 1. Chemical composition of bioactive glasses.

	SiO2	CaO	Na2O	P2O5	K2O	MgO	B2O3	Reference
45S5 Bioglass®	45 wt.%	24.5 wt.%	24.5 wt.%	6 wt.%	-	-	-	[53,54,55,56,57,58,59]
58S	58 wt.%	33 wt.%	-	9 wt.%	-	-	-	[114,115,116,117]
77S	77 wt.%	14 wt.%	–	9 wt.%	-	-	-	[118,119,120,121]
S53P4	53 wt.%	20 wt.%	23 wt.%	4 wt.%	-	-	-	[122,123,124]
13-93	53 wt.%	20 wt.%	6 wt.%	4 wt.%	12 wt.%	5 wt.%	-	[125,126,127]
13-93B1	34 wt.%	20 wt.%	6 wt.%	4 wt.%	12 wt.%	5 wt.%	19 wt.%	[128,129]
13-93B3	-	20 wt.%	6 wt.%	4 wt.%	12 wt.%	5 wt.%	53 wt.%	[129,130]
70S30C	71 wt.%	29 wt.%	-	-	-	-	-	[131,132,133,134]

). The second section describes the different stages of the bioactivity process occurring when a coated bone implant is immersed in a physiological environment. The third section provides a detailed description of the mechanisms involved in the electrophoretic deposition process. The last section reviews the current developments in the process of improving the properties of bioactive glass coatings by adding biocompatible elements to the glassy structure.

2. Bioactive Glasses

2.1. Bioglass^®

Larry Hench’s 45S5 Bioglass^® was the first bioactive glass produced in the history of biomaterials. This compound was synthesized for the first time in 1969 using the melt-quenching technique at temperatures above 1300 °C [53,54,55]. The 45S5 Bioglass^® chemical composition was obtained by mixing 45 wt.% SiO₂, 24.5 wt.% CaO, 24.5 wt.% Na₂O, and 6 wt.% P₂O₅. In the 1990s, an alternative method was developed to produce 45S5 Bioglass^® at low temperatures, the sol–gel process [135,136,137,138,139,140,141,142]. This synthesis method uses a solution of alkoxide precursors, such as tetraethyl orthosilicate (TEOS), typically used to obtain SiO₂. Under appropriate pH conditions, the solution is converted to a sol (colloidal solution) and then to a gel (solid network expended by a liquid phase) via hydrolysis and polycondensation reactions, respectively. 45S5 Bioglass^® powder is obtained by drying the gel. Sol–gel 45S5 Bioglass^® has a higher porosity and surface area, whereas melt-quenched 45S5 Bioglass^® has higher mechanical properties.

The chemical composition of Larry Hench’s 45S5 Bioglass^® is still the gold standard today. The first in vivo experiments after the discovery showed strong bonding to rat femurs after 6 weeks of implantation. The implanted pieces of 45S5 Bioglass^® had become impossible to move or remove from the rat bone [56,57,58,59]. These experimental results were rapidly confirmed in other animals, showing the ability of 45S5 Bioglass^® to bond both to hard and soft tissues [47,143,144]. Several concentrations of the four oxides were studied, but their bonding abilities to bone and soft tissues were limited at specific regions represented in Hench’s diagram (Figure 2) [145,146,147]. The diagram gathers all the compositions of quaternary glasses containing different concentrations of SiO₂, CaO, and Na₂O, with the concentration of P₂O₅ being kept constant at 6 wt.%. The bone-bonding ability of bioactive glasses is observed in region A. The soft tissue bonding ability of bioactive glasses is observed in region S. The composition of 45S5 Bioglass^® corresponds to region E. Glasses produced with compositions in regions B, C, and D do not bond to bone [128,148,149].

Larry Hench’s 45S5 Bioglass^® is a phosphosilicate glass in which SiO₂ and P₂O₅ are network formers and CaO and Na₂O are network modifiers (Figure 3). Each silicon atom of the glassy network is covalently bonded to four oxygen atoms in a tetrahedron structure [150,151]. The whole SiO₂ structure is disrupted by CaO and Na₂O, two network modifiers. They introduce non-bridging oxygen bonds that promote the dissolution of 45S5 Bioglass^® in aqueous environments [152,153,154]. The network former P₂O₅ is not mandatory to make the glass bioactive, but it acts as a nucleation site for the crystallization of apatite during the bioactivity process (see Section 3).

2.2. Other Bioactive Glasses

In addition to Larry Hench’s 45S5 Bioglass^®, the compositions of other phosphosilicate glasses have been explored to evaluate their bioactive properties [155,156,157,158,159,160,161]. At a higher SiO₂ concentration, 58S is the second-most studied bioactive glass in academic research [114,115,116,117]. In comparison to 45S5 Bioglass^®, there is no Na₂O in 58S, meaning that CaO is the only oxide acting as a network modifier of the structure. Another major bioactive glass is S53P4, widely used clinically for more than three decades. S53P4 supports bone repair and simultaneously inhibits bacterial growth [122,123,124]. A more complex compound is the bioactive glass 13-93, which is made of six different oxides. In addition to the four main oxides used to produce 45S5 Bioglass^®, 13-93 includes K₂O and MgO, which are both network modifiers of the glassy structure [125,126,127]. Several articles also describe bioactive phosphosilicate glasses of less usual compositions such as 60S, 65S, 75S, 77S, 80S, and 91S [118,119,120,121].

The addition of boric oxide (B₂O₃) to bioactive glasses is also well documented in the literature [129]. B₂O₃ acts as a network former in the glassy structure as SiO₂ or P₂O₅. Several studies have shown that boron possesses anti-inflammatory properties and stimulates angiogenesis in the body environment. Angiogenesis is the physiological process of producing new blood vessels that supply nutrients to bone cells during bone tissue repair [162,163]. The most usual compositions of borate bioactive glasses are 13-93B1, 13-93B3, 45S5B5, and HB5 [130,164,165,166,167].

A simpler bioactive glass containing only two compounds is 70S30C [131,132,133,134]. This bioactive glass contains 70 mol.% SiO₂ as a network former of the glassy structure, and 30 mol.% CaO as a network modifier, corresponding to 71 wt.% SiO₂ and 29 wt.% CaO.

3. Bioactivity of Glasses

Discovering bioactive glasses in 1969, Larry Hench introduced the concept of bioactivity and defined it as follows: “A bioactive material is one that elicits a specific biological response at the interface of the material which results in the formation of a bond between the tissues and the material” [53,146]. The bone-bonding ability of bioactive glasses results from the formation of a hydroxycarbonate apatite (HCA) layer on the surface of the glass. The precipitation of this layer is triggered by local supersaturation due to ionic releases from the partial dissolution of the glass under physiological conditions. This chemical process occurs in five stages [47,145,147]:

• Stage 1: Rapid ion exchange between the glass network modifiers (${\mathrm{N}\mathrm{a}}^{+}$ and ${\mathrm{C}\mathrm{a}}^{2+}$) with ${\mathrm{H}}^{+}$ ions (or ${\mathrm{H}}_3{\mathrm{O}}^{+}$) from the solution leads to the hydrolysis of the silica groups and the creation of silanol ($\mathrm{S}\mathrm{i}–\mathrm{O}\mathrm{H}$) groups on the glass surface:

  $$\mathrm{S}\mathrm{i}–\mathrm{O}–{\mathrm{N}\mathrm{a}}^{+}+{\mathrm{H}}^{+}+{\mathrm{O}\mathrm{H}}^{-}\to \mathrm{S}\mathrm{i}–{\mathrm{O}\mathrm{H}}^{+}+{\mathrm{N}\mathrm{a}}^{+}\left(\mathrm{a}\mathrm{q}.\right)+{\mathrm{O}\mathrm{H}}^{-}$$

  (1)
  The pH of the solution increases due to the consumption of ${\mathrm{H}}^{+}$ ions.
• Stage 2: The increase in the pH (or ${\mathrm{O}\mathrm{H}}^{-}$ concentration) leads to the attack of the ${\mathrm{S}\mathrm{i}\mathrm{O}}_2$ glass network, the dissolution of silica, in the form of silicic acid ${\mathrm{S}\mathrm{i}\left(\mathrm{O}\mathrm{H}\right)}_4$ into the solution, and the continuous formation of Si–OH groups on the glass surface:

  $$\mathrm{S}\mathrm{i}–\mathrm{O}–\mathrm{S}\mathrm{i}+{\mathrm{H}}_2\mathrm{O}\to \mathrm{S}\mathrm{i}–\mathrm{O}\mathrm{H}+\mathrm{O}\mathrm{H}–\mathrm{S}\mathrm{i}$$

  (2)
• Stage 3: Condensation and repolymerization of an amorphous ${\mathrm{S}\mathrm{i}\mathrm{O}}_2$-rich layer (typically 1 or 2 μm thick) occur on the surface of the glass depleted in ${\mathrm{N}\mathrm{a}}^{+}$ and ${\mathrm{C}\mathrm{a}}^{2+}$ by leaching:

  $$\mathrm{S}\mathrm{i}–\mathrm{O}\mathrm{H}+\mathrm{O}\mathrm{H}–\mathrm{S}\mathrm{i}\to \mathrm{S}\mathrm{i}–\mathrm{O}–\mathrm{S}\mathrm{i}+{\mathrm{H}}_2\mathrm{O}$$

  (3)
• Stage 4: The migration of ${\mathrm{C}\mathrm{a}}^{2+}$ and ${\mathrm{P}\mathrm{O}}_4^{3-}$ ions from the glass through the ${\mathrm{S}\mathrm{i}\mathrm{O}}_2$-rich layer and from the solution leads to the formation of an amorphous calcium phosphate (ACP) layer on the surface of the ${\mathrm{S}\mathrm{i}\mathrm{O}}_2$-rich layer.
• Stage 5: The amorphous calcium phosphate (APC) layer incorporates ${\mathrm{O}\mathrm{H}}^{-}$ and ${\mathrm{C}\mathrm{O}}_3^{2-}$ from the solution and crystallizes as an HCA layer.

Peitl et al. proposed the illustration in Figure 4 to schematically summarize the five chemical stages of the bioactivity process. This reaction sequence occurs when a bioactive phosphosilicate glass, like 45S5 Bioglass^®, is immersed in simulated body fluid (SBF) [168].

In addition to these five chemical stages resulting in the formation of the HCA layer, Larry Hench defined seven biological stages leading to the formation and growth of new bone tissue (Figure 5).

In stage 6, growth factors are adsorbed on the HCA layer where they activate M2 macrophages. Stage 7 corresponds to the action of M2 macrophages triggering the migration of mesenchymal stem cells and osteoprogenitor cells toward the bioactive glass surface, where they attach to the HCA layer (stage 8). Stem cells and osteoprogenitor cells differentiate in stage 9 to become osteogenic cells, which are precursors of osteoblast cells, like those present in bone tissue. The attachment and differentiation of osteoblasts generate extracellular matrix (ECM) components like type I collagen, enzymes, and glycoproteins (stage 10). Type I collagen is the most abundant protein in bones, representing more than 90% of the organic bone matrix [169].

In stage 11, the ECM is mineralized by hydroxyapatite nanocrystals, forming a three-dimensional network structure at the surface of the implant. The ECM provides structural and biochemical support to the surrounding bone cells, promoting their proliferation (stage 12). Bone growth is triggered at the implant surface, and the bioactivity process progressively continues degrading the bioactive glass coating, forming a greater ECM and supporting the development of new bone cells. The bioactivity process results in the osseointegration (or osteointegration) of the implant by creating an intimate connection with the surrounding bone tissue [170].

4. Electrophoretic Deposition (EPD)

Electrophoretic deposition of bioactive glass coatings on a metallic bone implant requires a stable colloidal suspension of bioactive glass particles in a solvent, typically water, ethanol, or a mixture of both [171,172]. When in contact with the solution, the surface of the particles becomes electrically charged due to electrostatic interactions with the ionic species of the solution. This surface charge maintains the stability of the colloidal suspension due to electrostatic interactions between the particles [173,174]. If two conductive electrodes connected to a generator are immersed in this colloidal suspension, the particles can be set in motion by applying an electric field between the electrodes. The particles move through the liquid in the direction of the oppositely charged electrode, where they progressively agglomerate to form a coating. If the surface of the particles is positively charged, they move toward and agglomerate on the cathode (cathodic EPD in Figure 6a). If the surface of the particles is negatively charged, they move toward and agglomerate on the anode (anodic EPD in Figure 6b).

After deposition, thermal annealing is necessary to evaporate the solvent and improve the cohesive and adhesive properties of the bioactive glass coating [175]. The experimental parameters influencing the electrophoretic deposition of bioactive glass coatings are the pH and stability of the colloidal suspension, the dielectric constant (ε) and viscosity (η) of the solvent, the average particle size, substrate conductivity, electric field and distance between the two electrodes, and the deposition time.

4.1. Surface Charge Formation

When bioactive glass powder is immersed into an aqueous solution, chemical interactions occur on the surface of each powder particle and induce the formation of a net surface charge [176,177]. The main surface phenomena involved in the process are the chemisorption of water molecules, the formation of silanol groups, protonation or deprotonation of the silanol groups, and ionic adsorption.

In an aqueous solution, chemisorbed water readily hydrolyzes the silica groups at the surface of the bioactive glass and creates silanol ($\mathrm{S}\mathrm{i}–\mathrm{O}\mathrm{H}$) groups (Figure 7) similarly to the descriptions of stages 1 and 2 of the bioactivity process (see Section 3).

Water molecules of the aqueous solution react with the silanol groups, resulting in either protonation (Reaction 4) or deprotonation (Reaction 5).

$$\mathrm{S}\mathrm{i}–\mathrm{O}\mathrm{H}+{\mathrm{H}}_2\mathrm{O}\to \mathrm{S}\mathrm{i}–{\mathrm{O}\mathrm{H}}_2^{+}+{\mathrm{O}\mathrm{H}}^{-}$$

(4)

$$\mathrm{S}\mathrm{i}–\mathrm{O}\mathrm{H}+{\mathrm{H}}_2\mathrm{O}\to \mathrm{S}\mathrm{i}–{\mathrm{O}}^{-}+{\mathrm{H}}_3{\mathrm{O}}^{+}$$

(5)

Lowe et al. deeply described the acid-base dissociation mechanisms and energetics at the silica-water interface [178]. They used ab initio calculations to demonstrate that the deprotonation of the silanol groups (Reaction 5) is the main reaction at pH values above 2. The protonation of the silanol groups according to Reaction 4 occurs only at a very low pH value (below 2), which is an unusual experimental condition for electrophoretic deposition. Consequently, in regular experimental conditions, the surface of a bioactive glass powder immersed in water becomes negatively charged due to the deprotonation of silanol groups.

This negatively charged surface influences the ionic distribution in the vicinity of the powder particle immersed in a polar medium like water. Positive ions (cations) of the solution are attracted toward the particle surface and remain attached to it (Figure 8).

These ions form the Stern layer of the electrical double-layer model in which positive counterions remain immobile on the particle surface [180]. The second layer of the double-layer model is called the “diffuse layer” and contains cations and anions in electrostatic interaction with the whole system. The diffuse layer ends with the shear slipping plane, the border with the bulk solution from which the charged particle has no more electrostatic influence. The electrical potential at this slipping plane is called the zeta potential (ζ), which is characteristic of the strength of all the charges carried by the particle [181].

4.2. Suspension Stability

The stability of a colloidal suspension depends on the strength of the electrostatic repulsions between the particles. The value of the zeta potential (ζ) must be high enough to prevent coagulation or flocculation of the particles [182]. Zeta potential values higher than +30 mV or lower than −30 mV are generally considered appropriate to obtain a stable suspension [183]. However, the zeta potential of the particles is highly influenced by the pH of the colloidal suspension (Figure 9). The highest values are observed at a low or high pH, whereas the value is zero at the isoelectric point, where the stability of the suspension is at the lowest.

4.3. Deposition Mechanism of Bioactive Glass Coating

Thanks to their zeta potential (ζ), the bioactive glass particles of the colloidal suspension can be set in motion by the electric field generated between two conductive electrodes. The particles move toward the oppositely charged electrode [184]. Under the influence of the electric field, the diffuse layer surrounding each bioactive glass particle is distorted and thinned (Figure 10). Some of the counterions of the diffuse layer leave the ion cloud because they are attracted by the other electrode in the opposite direction. Consequently, the zeta potential of each particle progressively decreases on their way toward the electrode surface, and the electrostatic repulsions between the particles are reduced [185]. Reaching the electrode at a high speed, the electric charges of the particles are neutralized, and they progressively accumulate via coagulation to form a uniform bioactive glass coating.

Zhang et al. proposed the following relation to calculate the deposition rate of a coating via electrophoretic deposition [187]:

$$\mathrm{w}={\displaystyle \frac{{\mathrm{m}}_0\hspace{0.17em}{\mathrm{\mu }}_{\mathrm{e}}\mathrm{S}\mathrm{E}}{\mathrm{V}}}\times {\mathrm{e}}^{({\displaystyle \frac{-{\mathsf{\mu }}_{\mathrm{e}}\mathrm{S}\mathrm{E}}{\mathrm{V}}}\mathrm{t})}$$

where w (kg·s⁻¹) is the deposition rate; m₀ (kg) is the initial mass of the powder in the suspension; V (m³) is the volume of the colloidal suspension; μ_e (m²·V⁻¹·s⁻¹) is the electrophoretic mobility of the particles; S (m²) is the surface of the coating; and E (V·m⁻¹) is the electric field between the two electrodes.

5. Current Developments in the Process and Perspectives

The most recent developments in the process explore solutions to improve the properties of bioactive glass coatings by adding biocompatible elements to the glassy structure. Several strategies are studied, including ionic substitution, deposition of composite coatings, and drug loading.

5.1. Ionic Substitution

As described in Section 2, the most common elements used to produce bioactive glasses are silicon, phosphorus, calcium, and sodium. Their biological properties can be enhanced by adding other ions to the glassy structure of the powder used in the electrophoretic deposition process. The purpose is to take advantage of the dissolution stage of the bioactivity process (see Section 3) to release biologically active ions into the physiological environment after implantation. The most usual ionic substitutions described in the literature involve potassium (K⁺), magnesium (Mg²⁺), zinc (Zn²⁺), silver (Ag⁺), strontium (Sr²⁺), fluorine (F⁻), cobalt (Co²⁺), copper (Cu²⁺), and boron (B³⁺). These ionic substitutions aim to provide osteogenic, angiogenic, anti-inflammatory, or antibacterial properties to the bioactive glass material [188,189]. The quantity of substituting element is generally low, typically a few percent of the whole mixture. Multi-substitution can be used to cumulate several biological enhancements of the bone implant [190,191]. More recently, less common elements have been used to substitute bioactive glass materials. They have been comprehensively reviewed by Pantulap et al., who thoroughly described the effect of these ions on the biological and physical properties of bioactive glasses [192]. They enhance hard and soft tissue repairs, and some of them provide extra functionalities such as anti-inflammatory, antibacterial, or anticancer properties, fluorescence, luminescence, and radiation shielding. Examples of substituting ions and their biological, physical, or chemical effects are reported in

Table 2. Ionic substitution in bioactive glasses.

Ions	Biological/Physical Effect	References
Monovalent Ions
${\mathrm{A}\mathrm{g}}^{+}$	Antibacterial activity	[193,194,195,196]
${\mathrm{C}\mathrm{l}}^{-}$	Osteogenesis	[197,198,199]
${\mathrm{F}}^{-}$	Osteogenesis	[200,201,202]
${\mathrm{K}}^{+}$	Osteogenesis	[203,204,205]
${\mathrm{L}\mathrm{i}}^{+}$	Antibacterial/osteogenesis	[206,207,208,209]
${\mathrm{R}\mathrm{b}}^{+}$	Antibacterial/drug carrier/osteogenesis	[210,211]
Divalent Ions
${\mathrm{B}\mathrm{a}}^{2+}$	Osteogenesis/radiation shielding	[212,213,214,215]
${\mathrm{C}\mathrm{o}}^{2+}$	Angiogenesis	[216,217,218,219]
${\mathrm{C}\mathrm{u}}^{2+}$	Antibacterial activity	[220,221,222,223]
${\mathrm{G}\mathrm{e}}^{2+}$	Osteogenesis/radiation shielding	[224,225]
${\mathrm{M}\mathrm{g}}^{2+}$	Osteogenesis	[226,227,228,229]
${\mathrm{M}\mathrm{n}}^{2+}$	Antibacterial/osteogenesis	[230,231,232,233]
${\mathrm{N}\mathrm{i}}^{2+}$	Drug carrier/osteogenesis/radiation shielding	[234,235,236]
${\mathrm{S}\mathrm{r}}^{2+}$	Osteogenesis	[237,238,239,240,241]
${\mathrm{Z}\mathrm{n}}^{2+}$	Antibacterial/anti-inflammatory/osteogenesis	[242,243,244,245]
Trivalent Ions
${\mathrm{B}\mathrm{i}}^{3+}$	Antibacterial/anticancer/osteogenesis	[246,247,248]
${\mathrm{C}\mathrm{e}}^{3+}$	Osteogenesis/antibacterial	[249,250,251,252]
${\mathrm{E}\mathrm{r}}^{3+}$	Osteogenesis/photoluminescence	[253,254]
${\mathrm{E}\mathrm{u}}^{3+}$	Drug carrier/photoluminescence	[255,256,257]
${\mathrm{F}\mathrm{e}}^{3+}$	Antibacterial/anticancer/osteogenesis	[258,259,260,261]
${\mathrm{G}\mathrm{a}}^{3+}$	Antibacterial/anticancer/osteogenesis	[262,263,264,265]
${\mathrm{H}\mathrm{o}}^{3+}$	Brachytherapy/osteogenesis	[266,267]
${\mathrm{L}\mathrm{a}}^{3+}$	Osteogenesis	[268,269,270]
${\mathrm{N}}^{3-}$	Osteogenesis	[271,272,273]
${\mathrm{S}\mathrm{m}}^{3+}$	Drug carrier/osteogenesis/photoluminescence	[274,275]
${\mathrm{T}\mathrm{b}}^{3+}$	Osteogenesis/photoluminescence	[276,277]
${\mathrm{Y}}^{3+}$	Anticancer/osteogenesis	[278,279,280]
${\mathrm{Y}\mathrm{b}}^{3+}$	Osteogenesis/photoluminescence	[254,281]
Tetravalent Ions
${\mathrm{S}\mathrm{e}}^{4+}$	Anticancer/drug carrier/osteogenesis	[282,283,284]
${\mathrm{T}\mathrm{e}}^{4+}$	Antibacterial/anticancer/osteogenesis	[285,286]
${\mathrm{Z}\mathrm{r}}^{4+}$	Antibacterial/osteogenesis	[287,288,289]
Pentavalent Ions
${\mathrm{N}\mathrm{b}}^{5+}$	Osteogenesis	[290,291,292,293]
${\mathrm{T}\mathrm{a}}^{5+}$	Antibacterial/osteogenesis	[294,295,296,297]
${\mathrm{V}}^{5+}$	Osteogenesis/photoluminescence/radiation shielding	[298,299,300]
Hexavalent Ions
${\mathrm{M}\mathrm{o}}^{6+}$	Drug carrier/osteogenesis	[301,302]

.

5.2. Composite Coatings

Electrophoretic deposition is a very versatile method. Any powder material added to the colloidal suspension can be simultaneously deposited with bioactive glass to produce a composite coating with improved properties [303]. The only requirement is the sign of all the zeta potentials be the same to move all the materials in the same direction and to deposit them on the same electrode. Similar zeta potential values are necessary to maintain the stability of the colloidal suspension. The main materials used to deposit composite coatings with bioactive glasses are bioceramics, polymers, and carbon nanotubes.

5.2.1. Bioactive Glass and Bioceramics

Several studies have reported on the simultaneous electrophoretic deposition of bioceramics with bioactive glasses. Bioceramics are biocompatible materials that are accepted by the human body without producing any adverse effect [39,40,41]. Some of them are bioinert, meaning they do not trigger any biological or chemical reaction with the surrounding environment [45,46]. They are typically combined with bioactive glass to improve the mechanical properties of the coating, such as its hardness, Young’s modulus, fracture toughness, bending strength, compressive strength, and wear resistance [304]. The most common bioinert ceramics used in composite coatings are Al₂O₃ [305], TiO₂ [306], ZrO₂ [307], ZnO [308], and Si₃N₄ [309]. Other bioceramics are bioactive or bioresorbable materials, meaning they trigger chemical and biological reactions in the physiological environment that promote bone growth and repair [46,47,48]. The main bioactive or bioresorbable bioceramics are calcium phosphates (hydroxyapatite, β-tricalcium phosphate, octacalium phosphate, and dicalcium phosphate dihydrate) [310], wollastonite (CaSiO₃) [311], calcium sulfate (CaSO₄) [312], and calcium carbonate (CaCO₃) [313]. Their bioactive behavior is linked to their solubility in physiological solutions, which is typically lower than the bioactivity of bioactive glasses. The kinetics of the bioactivity process can be fully controlled by combining two bioactive materials of different solubilities.

5.2.2. Bioactive Glass and Polymers

Polymers are used to improve the mechanical properties of bioactive glass coatings, specifically their hardness and adhesion to the bone implant [314]. They make the coating composition and structure closer to that of bone tissue, which is a composite material of collagen and apatite nanocrystals (calcium phosphate) [5,303]. Typical biocompatible polymers used to produce composite coatings are Polyether ether ketone (PEEK) [315], Poly(lactic-co-glycolic acid) (PLGA) [316], Polycaprolactone (PCL) [317], Poly(vinyl alcohol) (PVA) [318], Polymethyl methacrylate (PMMA) [319], and Poly(styrene-alt-maleic acid) (PSM) [320]. Natural biopolymers such as collagen [321], hyaluronic acid [322], alginate [323], chondroitin sulfate [324], zein [325], and chitosan [326] have also been described to produce composite coatings with improved mechanical and biological properties.

5.2.3. Bioactive Glass and Carbon Nanotubes (CNTs)

Carbon nanotubes are biocompatible, and they have high mechanical strength and flexibility [327]. They are combined with bioactive glass in the electrophoretic deposition process to produce composite coatings with improved mechanical properties and bioactivity [328,329,330].

5.3. Drug-Loaded Bioactive Glass Coatings

Drugs can be incorporated into bioactive glass coatings produced via electrophoretic deposition to achieve specific biological effects. The partial dissolution of the coating during the bioactivity process can be used to release active substances in the body after implantation. Drugs providing anti-inflammatory and antibacterial properties are mainly used. Several research studies succeeded in adding heparin [331], gentamicin [332], vancomycin [333], ibuprofen [334], ferulic acid [335], tetracycline hydrochloride [336], doxycycline [337], ampicillin [338], or lawsone [339].

6. Conclusions

This article reviewed the scientific literature dealing the electrophoretic deposition of bioactive glass coatings on metallic substrates to produce bone implants. After chemical and structural descriptions of Larry Hench’s 45S5 Bioglass^® and other bioactive glasses, the different stages of the bioactivity process were explained. According to the model proposed by Larry Hench, bioactive glass coatings immersed in a physiological environment trigger five chemical stages followed by seven biological stages, resulting in the formation and growth of new bone tissue. Then, the mechanisms involved in the electrophoretic deposition process were detailed. The surface of bioactive glass particles of a colloidal suspension becomes electrically charged when it is in contact with the solution. Under the influence of an electric field between two electrodes, the particles are set in motion and move toward an oppositely charged electrode, where they progressively accumulate to form a coating. Finally, the last part of this article reviewed the current developments of the process. The most recent solutions to improve the properties of bioactive glass coatings include the addition of biocompatible elements to the glassy structure, including ionic substitution, deposition of composite coatings, and drug loading.