Micro-Arc Oxidation in Titanium and Its Alloys: Development and Potential of Implants

Abstract

Titanium (Ti) and its alloys are widely recognized as preferred materials for bone implants due to their superior mechanical properties. However, their natural surface bio-inertness can hinder effective tissue integration. To address this challenge, micro-arc oxidation (MAO) has emerged as an innovative electrochemical surface modification technique. Its benefits range from operational simplicity and cost-effectiveness to environmental compatibility and scalability. Furthermore, the distinctive MAO process yields a porous topography that bestows versatile functionalities for biological applications, encompassing osteogenesis, antibacterial, and anti-inflammatory properties. In this review, we undertake an examination of the underlying mechanism governing the MAO process, scrutinize the multifaceted influence of various factors on coating performance, conduct an extensive analysis of the development of diverse biological functionalities conferred by MAO coatings, and discuss the practical application of MAO in implants. Finally, we provide insights into the limitations and potential pathways for further development of this technology in the field of bone implantation.

1. Introduction

The latest statistics show that by mid-century, the global population aged 65 or over is expected to more than double to a staggering 1.6 billion people [1]. An aging population means a larger population base of chronic diseases, and bone and joint diseases. In dentistry and orthopedics, implants, with titanium (Ti) and its alloys as the preferred choice among various materials, have become a common approach for treating conditions due to their exceptional biocompatibility and mechanical properties [2]. Despite the vast potential for Ti and its alloys, there are still limitations to surface bioactivity. They often exhibit biological inertness, limiting the close integration of implants with surrounding tissues. Furthermore, Ti alloys are susceptible to microorganisms, increasing the risk of postoperative infections, and threatening the prognosis of patients. To tackle these issues, apart from alloying [3,4,5], surface treatment proves to be an effective approach as well. Common methodologies encompass atmospheric plasma spraying [6], laser treatment [7], anodizing [8], micro-arc oxidation (MAO), and biomimetic deposition of apatite [9] or other biomaterials [10]. Among the various surface modifications available, MAO has garnered significant attention due to its simplicity of operation, cost-effectiveness, and applicability to complex devices. Through MAO, the surface of Ti-based implants can undergo alterations in microstructural features and chemical composition, rendering them more suitable for tissue adhesion and effectively reducing the risk of infections [11]. The specific impacts of MAO on the substrate surface can be found in

Table 1. A comparison table with the improved properties of MAO coatings on titanium alloys and uncoated ones.

Performance	Uncoated Ti Alloy	Coated Ti Alloy	Reference
Topography	Smooth or machined	Rough and porous	[12,13]
Chemical composition	Ti alloy	Ceramics dominated by TiO2	[12,13]
Hardness	Relatively not hard	Relatively hard	[13,14]
Wear resistance	Relatively not wear-resistant	Relatively wear-resistant	[15,16]
Corrosion resistance	Corrosion-resistant	More corrosion-resistant	[12,14,17]
Bioactivity	Biological inert	Biological activity	[12,18]

.

In this review, we aim to provide an overview of MAO applied to the preparation of coatings on Ti and its alloys for biomedical applications over the past 5 years, especially orthopedic and dental implants. This review encompasses an analysis of the factors influencing the coating preparation process, and an examination of the biological functionalities of these coatings and their applications. Finally, we highlight the challenges and future directions for the application of MAO in metal implants and make a vision for the development of surface coatings for metal implants.

2. The Process of MAO

2.1. The Mechanism of MAO

MAO involves plasma discharge, but it differs from traditional plasma electrolytic discharge techniques like plasma cutting or welding. It is often confused with anodic oxidation (AO), but they do indeed represent different processes and technologies. MAO typically requires a higher voltage and shorter processing times, and the type of electrolyte used, as well as the characteristics of the resulting coating, are significantly different from AO [19].

At present, most researchers divide the MAO process into three, which are spark discharge, micro-arc discharge, and arc discharge. Some add the early short anodizing phase, for a total of four stages [11,20].

The formation mechanism of MAO coatings is essentially similar, regardless of the electrolyte or control mode used. Taking the constant current mode as an example, Figure 1 illustrates the changes during the MAO process. In the first stage, as the current increases, the voltage linearly increases, initiating the traditional anodic oxidation phase, involving only electrochemical reactions. At this stage, a porous insulating film with a columnar structure perpendicular to the substrate is formed on the metal substrate, accompanied by gas generation, serving as precursors for igniting plasma discharge [21].

With the increase in current, the voltage on the coating also rises, moving on to the second and third stages. Once the voltage exceeds the critical threshold, known as breakdown voltage, it leads to spark discharge. There are various areas of defects, stress concentrations, and uneven thickness on the coating, and these vulnerable areas are the first to break down. This is the micro-arc discharge stage. In the later stage, when the maximum thickness of the porous oxide film is approached, the oxide film grows slowly with the increase in time, which is the arc discharge stage [23].

The phenomenon of micro-arc discharge and arc discharge is very similar, with only a slight difference in energy level. Two models of arc discharge have been proposed: one involving dielectric breakdown of the oxide film [24], and the other involving gas breakdown within micro-pores [25]. Both models result in the generation of high-energy plasma, possibly coexisting. Optical Emission Spectroscopy (OES) and Intensified Charge Coupled Device (ICCD)-based spectral systems provide real-time imaging of the MAO discharge process (Figure 2). The bright spots in the figure represent sparks, and over time, the intensity and range of sparks increase, indicating a larger pore size formed on the surface of the oxide film [26].

Stojadinovic et al. [27] used OES-ICDD to record real-time discharges. Two types of plasma micro-discharge were observed during this process, with electron densities of N_e = 3.8 × 10¹⁵ cm⁻³ and N_e = 4.5 × 10¹⁶ cm⁻³, and electron temperatures ranging from T_e = 3700 ± 500 K. Plasma discharge generates high temperatures, causing thermal decomposition of water molecules, producing more hydrogen and oxygen, further promoting plasma discharge. Oxides undergo melting ejections at high temperatures, encountering rapid solidification and recrystallization in contact with the cold electrolyte, depositing on the coating surface. A study [28] classified discharges into types a, b, and c, representing surface, internal, and intermediate discharges in coatings, leading to inward/outward growth of the coating. In this process, numerous white sparks scatter randomly and flicker on the workpiece surface [29]. As the discharge intensifies, the voltage increase rate slows down. With the melting and solidification of oxides, the thickness of coatings rapidly increases, and gas release and breakdown discharge generate pores. The number of sparks decreases, and the color changes from white to yellow. This marks the transition to the micro-arc discharge stage, a crucial phase in coating growth. As the reaction continues, the voltage stabilizes, entering the strong arc discharge stage. The spark discharge diameter gradually increases, turning orange-red, accompanied by a sharp noise. Intense discharge can cause spattering of coating materials and localized severe ablation characteristics.

Stress-induced cracking cannot be avoided during the rapid solidification process in contact with the cold electrolyte [30]. The resulting coating usually consists of a thinner internal dense layer and a thicker external porous layer, with thickness ranging from 1 to 100 μm [31]. To produce high-quality coatings, precise control of MAO process parameters is crucial, mainly including electrolytes, electrical parameters, and substrates.

2.2. Electrolyte

An electrolyte plays a pivotal role in the quality of MAO coatings. It directly influences the composition of coatings and exerts a significant impact on their growth mechanisms [31]. In the realm of MAO, electrolyte systems are traditionally classified by the anion type employed, including phosphate, silicate, borate, and aluminate. In recent years, organic phytic acid has gained recognition as an electrolyte for the MAO of biomedical alloys. Earlier practices involved acidic electrolytes [27], but recent mainstream electrolytes are typically alkaline with an ideal pH of around 13 [11]. The concentration, pH, electrical conductivity, additives, and temperature variations all influence coating formation.

In terms of electrolyte types, silicate and phosphate demonstrate better biocompatibility because Si and P elements promote osteogenesis [32]. When examining coating characteristics under similar conditions, it is observed that in the silicate electrolyte system, coatings grow more rapidly and to a greater thickness compared to other electrolyte systems. These coatings also exhibit increased roughness but relatively weaker adhesion to the metal substrates [33]. Consequently, such coatings often possess inferior wear resistance and corrosion resistance. The resulting bioceramic coatings primarily consist of rutile, anatase, and amorphous silica phases. In the phosphate electrolyte system, MAO coatings tend to be thinner but more compact, with lower roughness. They feature uniform, micrometer-sized pores, and exhibit strong adhesion to the substrate. These coatings consist primarily of anatase and rutile phases, providing excellent corrosion and wear resistance [30,34]. Silicate results in outward coating growth with poor adhesion due to oxide deposition, while phosphate promotes inward growth with strong adhesion linked to substrate oxidation. Combining silicate and phosphate achieves a balanced effect. The addition of calcium to the phosphate electrolyte can induce the formation of bone-like hydroxyapatite (HA), a biologically active material [32].

Due to the susceptibility of borates to hydrolysis, their solutions exhibit higher alkalinity and electrical conductivity. In the sodium tetraborate electrolyte, surfaces of Ti-based alloys undergo more intense discharges, resulting in the formation of a distinctive “channel-like” or “cortex-like” micro/nanostructure in the coatings (Figure 3). These coatings exhibit enhanced hydrophilicity and osseointegration. They are primarily composed of anatase, rutile TiO₂, and amorphous boron oxide [35]. Wang et al. [36] used sodium tetraborate to create a micro/nanoscale porous structure on the surface of a Ti-20Zr-10Nb-4Ta alloy. Results indicated that discharges in sodium tetraborate electrolytes lasted longer, compared to phosphate electrolytes, resulting in a larger pore size and higher porosity in the coatings.

When utilizing aluminate electrolytes, the coatings tend to precipitate aluminum oxide (or crystalline aluminum titanate, TiAl₂O₅), which provides higher hardness and stability. Consequently, these coatings often display excellent resistance to wear and corrosion [37,38]. Jiang et al. [39] employed phosphate, silicate, and aluminate electrolytes to prepare MAO coatings. They found that coatings produced in aluminate electrolytes had the densest structure. It is imperative to acknowledge that aluminate, which exhibits relatively poor stability, is susceptible to hydrolysis, leading to the formation of aluminum hydroxide. The presence of aluminum ions may potentially have detrimental effects on the human body. Consequently, the use of aluminate electrolytes in surface modification for biomedical metals has seen a gradual decline. In recent years, organic phytic acid has begun to find applications in the electrolytes of MAO [40,41]. It is an organic molecule rich in phosphorus elements, resulting in MAO films composed of anatase, rutile, and Ti₂O₇. It imparts outstanding biocompatibility to the coatings.

The anions in the electrolyte are involved in the discharge and coating formation processes, while cations typically influence the conductivity through concentration [42,43]. Conductivity is generally in the range of 5–100 mS/cm for the MAO electrolyte [41]. During the process, the conductivity of the electrolyte also changes, and the aging of the electrolyte is a noteworthy issue [44]. The conductivity is critical, influencing current and voltage within the narrow pores, thereby affecting the initiation of plasma discharge [27]. In the same anion electrolyte system, increasing conductivity leads to thicker coatings and reduced defects [45].

Currently, the incorporation of nano- and micro-particles into coatings is recognized as a primary approach for enhancing coating properties and expanding the scope of their chemical composition [46]. So far, various nanoparticles (NPs) (5–100 nm) and microparticles (0.18–44 μm) have been employed. These enhancements encompass a wide range of properties, including corrosion resistance, mechanical strength, tribological performance, and biocompatibility. For instance, the incorporation of graphene oxide (GO), TiO₂, and ZrO₂ enhances the corrosion resistance and tribological characteristics of coatings, while the inclusion of HA improves the biocompatibility [47]. In terms of composition, some elements are beneficial for improving the biological activity of the material and further enhancing the osseointegration, such as P [48], Si [49], Ca [50], Zr [6], and Sr [32]. The addition of Ag, Cu, and Zn imparts antibacterial properties to the coatings [51]. Incorporating particles into coatings is complex due to their limited solubility in aqueous solutions. These particles may remain unchanged in size and shape or melt during micro-discharge, influencing the composition, microstructure, and other properties of coatings. Key factors affecting these reactions at the solid–liquid interface include zeta potential, melting point, and particle size [22].

Zeta potential signifies the electrical charge at the particle–fluid interface, with high, negative zeta potential being favorable in electrolytes for stability, resisting agglomeration, and settling [15]. Adjusting the pH of the electrolyte can shift the zeta potential toward negativity. By measuring the zeta potential, Necula et al. [52] established optimal conditions for Ag incorporation in the Keronite electrolyte. Melting point and particle size are two other critical factors affecting incorporation, because of the elevated local temperatures during the MAO treatment (~2700 °C) [53]. Particles with high melting points, such as CeO₂ and SiC, tend to remain inert or partially reactive regardless of their size. On the other hand, particles with lower melting points are more prone to reactive incorporation, resulting in mixed-oxide coatings within the discharge channels during the MAO process. As an example, SiO₂ particles with a low melting point have been successfully incorporated into MAO coatings. Nanometric particles (around 12 nm) showed active integration, while micrometric particles (1–5 μm) remained inert [54]. To mitigate additive agglomeration, various strategies are available, including the utilization of chelating agents and organic acids, such as nitrilotriacetic acid (C₆H₉NO₆), phosphoric acid (H₃PO₄), phytic acid (C₆H₁₈O₂₄P₆), and disodium EDTA (Na₂EDAT·2H₂O). These approaches help create stable and chemically compatible solutions or suspensions [55].

Electrolyte temperature significantly impacts the discharge. There is a non-linear relationship between temperature and coating thickness. Initially, a higher temperature improves corrosion and wear resistance, peaking at 40 °C, but then it declines. Rapid temperature changes in the electrolyte pose challenges to process stability [56]. Electrolyte temperature fluctuations impact the properties of MAO coatings. Higher temperatures can lead to excessive discharging, causing uneven surface topography in the coating [11]. Coatings produced at 278 K have lower porosity than those at higher temperatures due to larger pores and increased pore density, resulting in reduced adhesion to substrates [57]. To obtain high-quality coatings, it is advisable to use a cooling system to keep the electrolyte temperature around 20 °C, with a maximum not exceeding 40 °C [58].

2.3. Electrical Parameters

In the MAO process, the performance of the coating is influenced not only by the electrolyte but also by various electrical parameters, including processing time, current mode, current density, and frequency.

In the standard power mode of the MAO process, different current modes can be employed, including direct current (DC) and alternating current (AC). Currently, the pulsed current is also widely used in MAO treatment, including both DC pulse power and AC pulse power, whether a unipolar or bipolar pulse [59]. The DC mode results in slower oxide growth and increased porosity, though it offers less control. Implementing pulsed DC may improve control over discharge duration and reduce energy consumption [60]. Compared to DC power, AC power can swiftly mitigate electrode polarization by intermittently interrupting the electric arc. It provides increased flexibility, improves process control efficiency, and enhances coating quality. Currently, the use of AC-pulsed bipolar MAO power sources is widespread due to their capacity to produce superior coatings. These coatings exhibit a greater thickness, higher micro-hardness, and enhanced adhesion to substrates [61]. This effectiveness is attributed to their ability to mitigate the impact of high-intensity plasma discharges and reduce the occurrence of high-temperature spikes [62]. In different power modes, constant current and constant voltage modes can be employed. A work [63] demonstrates that, when using the constant direct current mode, extending the duration at a low current density (2 A/dm²) helps prolong the micro-arc phase and results in the formation of a porous layer with fewer internal defects. The compact internal structure enhances the hardness and corrosion resistance. Conversely, with higher current density, the discharge enters the large arc phase earlier, leading to an increase in internal defects and a decrease in adhesion. Seo et al. [64] treated the surfaces of CP-Ti and Ti-6Al-4V ELI discs under pulsed DC with constant voltage conditions. Increasing voltage led to a larger grain size and surface pore size, and more rutile content, while anatase content decreased. A high voltage caused surface cracking. Both constant current and constant voltage modes have pros and cons. A constant current ensures uniform arc discharges but can generate excess heat, affecting coating uniformity. A constant voltage allows better control of arc discharges and reduces thermal effects, though it may slightly compromise coating uniformity.

Both unipolar and bipolar pulse modes introduce variable parameters like frequency, duty cycle, and anode-to-cathode current ratio, necessitating a deeper understanding of their impact on coatings. Previous research indicates that frequency plays a crucial role in coating performance. Lower frequencies result in rougher coatings with larger pores, making AC pulses more effective than DC pulses for coating formation. Higher frequencies lead to coatings with a higher density of finer pores [65]. Wang et al. [66] created Ca- and P-containing MAO films on Ti substrates at frequencies from 100 to 5000 Hz. Higher frequencies decrease crystallinity but increase Ca and P content. Films at higher frequencies have smaller pores on larger pore walls, improving film wettability.

Maintaining a constant voltage and frequency, an increased duty cycle lengthens the discharge period in a single-polarity pulse, resulting in thicker coatings, additional pores, and greater surface roughness [67,68]. However, excessively high duty cycles can damage the coating. Babaei et al. [69] prepared TiO₂-ZrO₂ composite coatings on CP-Ti, finding that different micro-discharge states influence the performance of coatings prepared under varying duty cycles. Higher duty cycles lead to an increased coating thickness and greater surface roughness. Pavarini et al. [70] created Cu and Zn co-doped MAO coatings with antimicrobial and osteogenic properties using pulsed DC in a sodium tetraborate electrolyte system. Higher frequencies (800 Hz) produced thinner, less uniform, and less porous coatings with smaller pores. Similar effects occurred with lower duty cycles (10%). Increasing the duty cycle and reducing the frequency resulted in a higher proportion of rutile in the oxide composition.

Pulsed bipolar MAO power sources are currently widely used for increasing coating thickness and ensuring strong adhesion to substrates. If the cathodic pulse is increased, or the ratio of cathodic current to anodic current (Ic/Ia) is increased [71], the residual discharge channels decrease, leading to higher coating density though the specific cathodic discharge mechanism remains to be elucidated [72]. Sun et al. [73] used α-Al₂O₃ to create a hard ceramic coating on a Ti-6Al-4V alloy. Lower Ic/Ia ratios led to thicker coatings, indicating higher anodic current density promotes growth. Conversely, higher Ic/Ia ratios resulted in thicker dense layers, reduced surface roughness, and increased coating hardness.

The duration of the MAO process is critical. In general, as the time is prolonged, the discharges become more intense, the number of micro-pores decreases, their size increases, and the coating thickness increases [29]. Wu et al. [74] found that as the processing time extended, the coating exhibited the mentioned phenomena, but some central pore sections closed and exhibited shrinkage. This is because the molten oxides flowing at the discharge locations sealed the pores.

MAO is an energy-intensive process, with actions like increasing current density, elevating voltage, reducing frequency, extending duty cycle, and prolonging processing time all equivalently increasing energy input over time [22]. Higher instant energy levels intensify discharges, resulting in more micro-pores in the coating, leading to increased roughness and a higher rutile phase content. Moderate energy increases create a complex porous structure, but excessive energy input may harm the coating. This can be addressed by raising the cathodic current, leading to the adoption of alternating current and pulse power sources. There is ongoing research into soft spark discharge states, which emit less heat, reducing excessive and damaging discharges [75]. This fosters uniform electrode reactions, lowers energy consumption, and promotes coatings rich in crystalline phases. However, soft spark discharge reduces surface porosity and roughness, and its potential benefits for enhancing bioactivity require further investigation [76].

The decision on when to use each mode or parameter requires a comprehensive evaluation of the desired coating characteristics, process control requirements, material and shape of the workpiece, energy efficiency, and application environment. The optimal choice often involves experimentation and optimization to meet specific project requirements.

2.4. Substrate

The composition, topography, and properties of the coating are significantly influenced by the composition and structure of the metal substrate. Additionally, pre-treatments conducted before MAO can also impact the grain size and surface topography of the substrate, subsequently affecting the topography and quality of the MAO coatings. Currently, commonly used Ti and its alloys in clinical applications include CP-Ti, Ti-6Al-4V, and Ti-6Al-7Nb, which have been the focus of substantial surface modification research. However, the rise of β-Ti alloys for hard tissue implants, owing to their superior biocompatibility and reduced elastic modulus for reduced “stress shielding”, has led to an increasing emphasis [77].

Chen et al. [78] utilized NaH₂PO₄ as an electrolyte for MAO treatment of a Ti-13Cr-3Al-1Fe alloy and found that the response to electrical parameters is akin to the trend observed in α and α + β alloys. Cimenoglu et al. [79] investigated MAO-treated Ti-6Al-4V and Ti-6Al-7Nb alloys using the same parameters. Both exhibited a similar thickness of approximately 10 μm. However, the coating on Ti-6Al-4V contained HA precipitates, while the Ti-6Al-7Nb coating had granular oxides with calcium titanate. In the study by Wang [80], an MAO process was employed under identical conditions on pure Ti, Ti-6Al-4V, and Ti-35Nb-2Ta-3Zr substrates. The results revealed distinct topographies of coatings on the three different substrates (Figure 4). Notably, the Ti-35Nb-2Ta-3Zr surface exhibited a porous topography with increased crystalline anatase content, which effectively enhanced the deposition of HA. Additionally, corrosion resistance and hydrophilicity tests indicated superior performance for the coating.

Material pretreatment processes alter the surface topography and grain size, consequently affecting the properties of coatings. Physical methods like surface shot peening, severe plastic deformation (SPD), and laser surface remelting (LSR) treatments are primarily considered. High-energy shot peening (HESP) enhances the wettability and surface-free energy of MAO coatings, significantly improving bioactivity on Ti-based substrates [49]. Equal channel angular pressing (ECAP) on Ti surfaces results in ultrafine grains and increases the thickness and porosity of MAO coatings. Furthermore, the dense microstructure enhances the corrosion resistance of MAO coatings [14]. LSR pretreatment, in combination with MAO processing, modulates coating topography. Increasing the energy density of LSR pretreatment initially decreases surface roughness before subsequently increasing it. This enhances coating densification and bond strength with the substrate, thereby improving corrosion resistance [81]. In conclusion, appropriate pretreatment can enhance the performance of MAO coatings. Additionally, subjecting alloy substrates to solution treatment, quenching, and varying aging treatments before MAO results in changes in coating pore size, surface roughness, and porosity with increasing aging time. This is likely due to enhanced substrate electrical conductivity from the aging treatment, leading to higher current density during the MAO process [82]. The comprehensive effects of electrolytes, electrical parameters, and substrates can be found in Table 2.

At present, additive manufacturing, also known as 3D printing, is progressively emerging as the preferred method for custom orthopedic implants. It allows the production of complex structures that surpass the capabilities of traditional techniques, all while preserving the physical and chemical advantages inherent to Ti and its alloys. Implants produced through this method offer an elevated level of patient-specific adaptability. In the context of employing MAO to enhance the surface bioactivity of 3D-printing Ti alloy implants, more detailed comparisons are available in other comprehensive reviews [2,83].

Obtaining the appropriate parameters requires multiple experiments. Researchers can start by choosing an experimentally suitable parameter range based on existing literature. They should then refine these parameters based on observed multifaceted effects of parameter changes on the coating. Due to the absence of a unified model, conducting multiple experiments is essential. The alteration of a single parameter in the MAO process can result in modifications across multiple coating properties, owing to the intricate interplay of various factors. Consequently, adjusting a specific parameter may necessitate the simultaneous coordination of other parameters to uphold the desired coating quality. This intricate balance poses a significant challenge in the production process of MAO coatings. Different applications of implants have varying requirements for the surface coating, leading to variations in the parameters used. Coating performances prepared under the same parameters with different devices may exhibit slight differences, making it challenging to define optimal parameters [22]. The parameters influencing the coatings are comprehensively addressed here to serve as a reference for subsequent researchers. After determining the parameters, real-time online monitoring systems can be introduced to track key parameters such as current, voltage, and electrolyte concentration. Through feedback control, maintaining parameters within a stable range ensures consistent and reproducible coating performance.

Table 2. Factors affecting coating performance.

Number	Substrate	Electrolyte	Electrical Regimes	Conclusion	Reference
1	Ti	Na3PO4, Na2SiO3, NaAlO2, composite electrolytes (Na3PO4 + Na2SiO3 + NaAlO2)	0.1–0.3 A/cm2 current densities	The pores of coatings made in Na3PO4 and composite electrolytes are smaller and more uniform than those in other electrolytes.	[39]
2	Ti	KH2PO4, Ca(OH)2 or Ca(HCOO)2, Na2(EDTA)	Three different fixed set current densities: 50, 100, and 150 mA/cm2	Porous coating promotes apatite formation and resists corrosion.	[55]
3	Ti	(CH3COO)2Ca·H2O and NaH2PO4·2H2O	Different voltages, currents, durations	HA-containing flower-like Structure Coatings have good biological activity.	[84]
4	Ti	Ca(OOCCH3)2, Ca(H2PO4)2, and Na2(EDTA)	Pulsed DC, duty cycle at 30%, 450 V, 5 min, 100–5000 Hz	Frequency affects the crystallinity, composition, topography, and wetting ability of the oxide film.	[66]
5	Ti	NaH2PO4, Na2ZrO3, and Na2SiO3 of different concentrations	Pulsed DC, 500 V, 1000 Hz, different duty cycle	Concentration and duty cycle affect topography and photocatalytic activity.	[69]
6	Ti (high-energy shot peening, HESP)	C3H7Na2O6P·5H2O, Ca (CH3COO)2·H2O, Na2SiO3·9H2O, and Cu(CH3COO)2· H2O	480 V, 5 min	Porous antimicrobial coatings are prepared.	[49]
7	Ti (equal channel angular pressing, ECAP)	NaH2PO2 and (CH3COO)2Ca	350 mA/cm2, 8 min	Porous microcrack coatings are prepared.	[14]
8	Ti	K3PO4 and KOH	Constant current density of 100 mA/cm2, frequency of 6 kHz, duty cycle at 50%, 6 min	Coating with pancake-like topography has corrosion resistance.	[46]
9	Ti	Different concentrations of KOH	160 V, 1 min	1 M KOH, Ca/P = 1.69. The higher the concentration of KOH, the stronger the corrosion resistance of the coating.	[42]
10	Ti	Na2B4O7·10H2O, Na2O2SiO2·2H2O, Cu(CH3COO)2, Zn(CH3COO)2, NaOH	AC, 300 V, different frequencies, and duty cycles	Duty cycle and frequency affect coating thickness, surface uniformity, and porosity.	[70]
11	Ti	Na2B4O7·10H2O	Pulsed DC, 465 V, 600 Hz, and 9%	“Cortex-like” micro/nanostructured coating has improved biocompatibility.	[35]
12	Ti, Nb, Mg, Al, Zr, Ta	NaOH and Na2SiO3	Unipolar positive or bipolar working pulse	Soft sparks are more obvious on the surfaces of Mg, Al, Zr, and Ta, while Ti and Nb have only a small amount of spark softening.	[85]
13	Ti-6Al-4V	Ca(CH3COO)2, Ca(C3H7O6P), Mn(CH3COO)2·4H2O, Mg(CH3COO)2·4H2O, Sr(CH3COO)2·0.5H2O, Zn(CH3COO)2, Na2SiO3	Pulsed DC, 280 V, 3 min	The coating surface is soaked in simulated body fluid (SBF) to generate bone-like apatite.	[32]
14	Ti-6Al-4V	Na3PO4, NaOH, Na2SiO3	Fixed DC voltage (270 V), 5 min	Coatings prepared in silicate-based electrolytes achieved the most uniform structure, with lower porosity.	[33]
15	Ti-6Al-4V	NaAlO2, NaF, KOH, CuSO4, Na2(EDTA)	Constant current of 2 A/cm2	It is porous at about 1μm. The addition of Cu ions makes the coating uniform and reduces the roughness.	[38]
16	Ti-6Al-4V	Na2HPO4, HA microns, and NPs	Constant potential of 250 V and then pulse unipolar or pulse bipolar constant current of 300 mA/cm2	Micron and nanoscale HA have different effects on the microstructure of the coating, but both improve the scratch resistance and bioactivity of the coating.	[86]
17	Ti-6Al-4V	Na2AlO2, Na3PO4	Bipolar pulse	With an appropriate increase in cathodic pulses, the coating density increases.	[72]
18	Ti-6Al-4V	Na2AlO2, Na3PO4	Bipolar pulse	As the Ic/Ia ratio increases, the density and hardness of the coating increase, and the roughness decreases.	[73]
19	Ti-6Al-4V	Na2SiO3, NaH2PO2, Na2MoO4	520 V, 50 Hz, 10%, 30 min	Corrosion resistance is improved.	[81]
20	Ti-20Zr-10Nb-4Ta	Na2B4O7·10H2O	DC power, 465 V, 600 Hz, and 9%	Hierarchical porous coatings have good biocompatibility.	[36]
21	Ti-15V-3Al-3Cr-3Sn	Different electrolyte temperatures of 278–313 K, K2Al2O4, Na3PO4, NaOH	square waveform with 2.0 kA/m2 and −1.0 kA/m2 at 100 Hz	Wear resistance is improved at low temperatures.	[87]
22	Ti-13Cr-3Al-1Fe	NaH2PO4, pH = 9.2	Different voltages and durations	As the discharge voltage and treatment time increase, the oxide layer thickens, and the pore size and surface roughness increase.	[78]
23	Ti6Al4V and Ti6Al7Nb	(CH3COO)2Ca·H2O and Na3PO4	Bipolar power, +500 V and −83 V, 5 min	Two alloy coatings have different characteristics under the same conditions.	[79]
24	Ti-29Nb-13Ta-4.6Zr(hot forge)	CaOH, Na3PO4·12H2O, pH = 12.03, conductivity = 13.20 mS/cm	AC bipolar, different frequencies, and duty cycles	Frequency and duty cycle affect form and thickness, which in turn affects corrosion resistance.	[68]
25	Ti-25Ta-10Zr-15Nb and Ti-25Ta-20Zr30Nb	Ca(C2H3O2)2, Mg(C2H3O2)2, Glycerophosphate	300 V, 2.5 A, 1 min	Both alloys exhibit porous structures with different pore sizes.	[88]
26	Ti, Ti-6Al-4V, and Ti-15V-3Al-3Cr-3Sn	K2Al2O4, Na3PO4, NaOH	Constant AC of 1.5 kA/m2 up to a maximum peak voltage of 400 V	The coating has excellent wear resistance.	[57]
27	Ti, Ti-6Al-4V, Ti35Nb-2Ta-3Zr	Na2SiO3·9H2O, (HOCH2CH2)3N, pH = 11.5	Voltage of +400 V, −50 V, duty cycle at +16%, −10%, 500 Hz	The coating of Ti-35Nb-2Ta-3Zr is durable and has good hydrophilicity.	[80]

Table 2. Factors affecting coating performance.

Number	Substrate	Electrolyte	Electrical Regimes	Conclusion	Reference
1	Ti	Na3PO4, Na2SiO3, NaAlO2, composite electrolytes (Na3PO4 + Na2SiO3 + NaAlO2)	0.1–0.3 A/cm2 current densities	The pores of coatings made in Na3PO4 and composite electrolytes are smaller and more uniform than those in other electrolytes.	[39]
2	Ti	KH2PO4, Ca(OH)2 or Ca(HCOO)2, Na2(EDTA)	Three different fixed set current densities: 50, 100, and 150 mA/cm2	Porous coating promotes apatite formation and resists corrosion.	[55]
3	Ti	(CH3COO)2Ca·H2O and NaH2PO4·2H2O	Different voltages, currents, durations	HA-containing flower-like Structure Coatings have good biological activity.	[84]
4	Ti	Ca(OOCCH3)2, Ca(H2PO4)2, and Na2(EDTA)	Pulsed DC, duty cycle at 30%, 450 V, 5 min, 100–5000 Hz	Frequency affects the crystallinity, composition, topography, and wetting ability of the oxide film.	[66]
5	Ti	NaH2PO4, Na2ZrO3, and Na2SiO3 of different concentrations	Pulsed DC, 500 V, 1000 Hz, different duty cycle	Concentration and duty cycle affect topography and photocatalytic activity.	[69]
6	Ti (high-energy shot peening, HESP)	C3H7Na2O6P·5H2O, Ca (CH3COO)2·H2O, Na2SiO3·9H2O, and Cu(CH3COO)2· H2O	480 V, 5 min	Porous antimicrobial coatings are prepared.	[49]
7	Ti (equal channel angular pressing, ECAP)	NaH2PO2 and (CH3COO)2Ca	350 mA/cm2, 8 min	Porous microcrack coatings are prepared.	[14]
8	Ti	K3PO4 and KOH	Constant current density of 100 mA/cm2, frequency of 6 kHz, duty cycle at 50%, 6 min	Coating with pancake-like topography has corrosion resistance.	[46]
9	Ti	Different concentrations of KOH	160 V, 1 min	1 M KOH, Ca/P = 1.69. The higher the concentration of KOH, the stronger the corrosion resistance of the coating.	[42]
10	Ti	Na2B4O7·10H2O, Na2O2SiO2·2H2O, Cu(CH3COO)2, Zn(CH3COO)2, NaOH	AC, 300 V, different frequencies, and duty cycles	Duty cycle and frequency affect coating thickness, surface uniformity, and porosity.	[70]
11	Ti	Na2B4O7·10H2O	Pulsed DC, 465 V, 600 Hz, and 9%	“Cortex-like” micro/nanostructured coating has improved biocompatibility.	[35]
12	Ti, Nb, Mg, Al, Zr, Ta	NaOH and Na2SiO3	Unipolar positive or bipolar working pulse	Soft sparks are more obvious on the surfaces of Mg, Al, Zr, and Ta, while Ti and Nb have only a small amount of spark softening.	[85]
13	Ti-6Al-4V	Ca(CH3COO)2, Ca(C3H7O6P), Mn(CH3COO)2·4H2O, Mg(CH3COO)2·4H2O, Sr(CH3COO)2·0.5H2O, Zn(CH3COO)2, Na2SiO3	Pulsed DC, 280 V, 3 min	The coating surface is soaked in simulated body fluid (SBF) to generate bone-like apatite.	[32]
14	Ti-6Al-4V	Na3PO4, NaOH, Na2SiO3	Fixed DC voltage (270 V), 5 min	Coatings prepared in silicate-based electrolytes achieved the most uniform structure, with lower porosity.	[33]
15	Ti-6Al-4V	NaAlO2, NaF, KOH, CuSO4, Na2(EDTA)	Constant current of 2 A/cm2	It is porous at about 1μm. The addition of Cu ions makes the coating uniform and reduces the roughness.	[38]
16	Ti-6Al-4V	Na2HPO4, HA microns, and NPs	Constant potential of 250 V and then pulse unipolar or pulse bipolar constant current of 300 mA/cm2	Micron and nanoscale HA have different effects on the microstructure of the coating, but both improve the scratch resistance and bioactivity of the coating.	[86]
17	Ti-6Al-4V	Na2AlO2, Na3PO4	Bipolar pulse	With an appropriate increase in cathodic pulses, the coating density increases.	[72]
18	Ti-6Al-4V	Na2AlO2, Na3PO4	Bipolar pulse	As the Ic/Ia ratio increases, the density and hardness of the coating increase, and the roughness decreases.	[73]
19	Ti-6Al-4V	Na2SiO3, NaH2PO2, Na2MoO4	520 V, 50 Hz, 10%, 30 min	Corrosion resistance is improved.	[81]
20	Ti-20Zr-10Nb-4Ta	Na2B4O7·10H2O	DC power, 465 V, 600 Hz, and 9%	Hierarchical porous coatings have good biocompatibility.	[36]
21	Ti-15V-3Al-3Cr-3Sn	Different electrolyte temperatures of 278–313 K, K2Al2O4, Na3PO4, NaOH	square waveform with 2.0 kA/m2 and −1.0 kA/m2 at 100 Hz	Wear resistance is improved at low temperatures.	[87]
22	Ti-13Cr-3Al-1Fe	NaH2PO4, pH = 9.2	Different voltages and durations	As the discharge voltage and treatment time increase, the oxide layer thickens, and the pore size and surface roughness increase.	[78]
23	Ti6Al4V and Ti6Al7Nb	(CH3COO)2Ca·H2O and Na3PO4	Bipolar power, +500 V and −83 V, 5 min	Two alloy coatings have different characteristics under the same conditions.	[79]
24	Ti-29Nb-13Ta-4.6Zr(hot forge)	CaOH, Na3PO4·12H2O, pH = 12.03, conductivity = 13.20 mS/cm	AC bipolar, different frequencies, and duty cycles	Frequency and duty cycle affect form and thickness, which in turn affects corrosion resistance.	[68]
25	Ti-25Ta-10Zr-15Nb and Ti-25Ta-20Zr30Nb	Ca(C2H3O2)2, Mg(C2H3O2)2, Glycerophosphate	300 V, 2.5 A, 1 min	Both alloys exhibit porous structures with different pore sizes.	[88]
26	Ti, Ti-6Al-4V, and Ti-15V-3Al-3Cr-3Sn	K2Al2O4, Na3PO4, NaOH	Constant AC of 1.5 kA/m2 up to a maximum peak voltage of 400 V	The coating has excellent wear resistance.	[57]
27	Ti, Ti-6Al-4V, Ti35Nb-2Ta-3Zr	Na2SiO3·9H2O, (HOCH2CH2)3N, pH = 11.5	Voltage of +400 V, −50 V, duty cycle at +16%, −10%, 500 Hz	The coating of Ti-35Nb-2Ta-3Zr is durable and has good hydrophilicity.	[80]

3. Surface Biologic Function of MAO

3.1. Osteointegration

When an implant is exposed to the internal environment, it selectively adsorbs various proteins, depending on the physical and chemical characteristics of its surface. Cells initiate their connection with the implant by interacting with these unique proteins [89]. Surface parameters of the implant, including surface topography, chemical composition, and charge, all contribute to regulating specific interactions between cells and biomaterials.

On smooth surfaces, BMSCs are more prone to differentiate into fibroblasts rather than osteoblasts (OBs). Additionally, fibroblasts also adhere more easily to smooth surfaces. On the other hand, bone cells tend to adhere, proliferate, and differentiate more readily on surfaces with rough nanoscale patterns (Figure 5). Surfaces with moderate roughness and porosity provide a larger bonding area and establish a strong mechanical interlock, promoting a more stable implant–bone integration. Osteoblasts cultured on such rough surfaces demonstrate increased collagen production and enhanced calcification processes [90,91].

The ideal implant surface should mimic natural bone, featuring mineralized large pores and nanoscale components. Hierarchical structure, from the nanoscale (collagen molecules and minerals) to microscale (bone units), guides cells’ behavior [92]. Research suggests that multiscale topographical features on the implant surface are linked to increased osteoblast differentiation [93].

MAO can modify the surfaces of metal implants to give them specific characteristics. It can create various topological structures on the substrate, and incorporate substances for angiogenesis or osteointegration into the coatings.

It is widely acknowledged that a rough, porous surface contributes to improved osteogenic activity. Nevertheless, there is currently no consensus regarding the optimal specific values for roughness, pore size, and pore geometry [94]. Yamasaki et al. [95] highlighted that the interconnected micropores in the range of 2–10 μm can bestow osteoinductive properties upon the scaffold. Shalabi et al. [96] found that bone tissue bonded effectively to rough Ti surfaces within the 1 to 100 μm range, with higher surface roughness providing a greater surface area. Currently, the surface topographies generated through MAO exhibit diversity, encompassing porous [38], hierarchical porous [97,98], and “cortex-like” (or “worm-like”) [35] structures. Li et al. [98] employed a two-step MAO process to create pores of different sizes. In the first step, large pores (100–300 μm) were generated using NaNO₃ and NaOH electrolytes. Subsequently, a second MAO treatment was conducted using a Na₂B₄O₇ electrolyte, resulting in the formation of micro-pores (3–10 μm) as well as sub-micron/nano-pores (80–200 nm). This innovative tri-level structure not only enhances hydrophilicity but also holds the potential to improve the adhesion of osteoblasts to the implant and enhance implant fixation. Moreover, they refined the electrolyte composition by introducing CaO and SrO into Na₂B₄O₇ [46]. Following a one-step MAO treatment, both Ca and Sr were integrated into the TiO₂ hierarchical porous coating. This hierarchical structure exhibits a high level of porosity and superhydrophilicity, and the incorporated Ca and Sr have shown superior efficacy in promoting the proliferation of human-bone-marrow-derived mesenchymal stem cells (hBMSCs). This highlights that, beyond surface topography, the chemical composition of MAO coatings plays a crucial role in influencing the biological activity of the coatings.

Currently, a highly researched area revolves around the development of biomimetic calcium phosphate (CaP) coatings on surfaces using electrolytes containing both Ca and P. Researchers are dedicated to optimizing the parameters that influence coating formation during the MAO process, with a specific emphasis on the composition of the electrolyte. They aim to achieve a Ca/P ratio that closely matches the natural bone value of 1.67 [48].

Typically, immersion in SBF is employed to assess the ability to form HA on its surface, thus quantifying its bioactivity. The general formula of HA is Ca₁₀(PO₄)₆(OH)₂. Herein, Ca²⁺ can be completely or partially substituted by Sr²⁺, Si⁴⁺, Mg²⁺, Mn²⁺, and Zn²⁺, and OH⁻ can be replaced by halide ions F⁻, Cl⁻, and Br⁻. Therefore, in addition to Ca and P, adding Si, Sr, Mn, Zn, and other elements to the electrolyte may also improve the osteogenic activity of the surface. Plekhova et al. [18] assessed the morphological and functional status of bone-marrow-derived mesenchymal stem cells (BMSCs) cultured on Ti-CaP coatings formed by MAO. There were no cytotoxic effects observed on the cultured cells. Notably, the high expression of receptors (CD90, CD29, and CD106) along with enhanced synthesis of osteocalcin and osteopontin, as well as observable changes in the surface structure of BMSCs adhered to the samples, collectively confirmed the osteoinductive properties of the calcium phosphate MAO coating. Park et al. [99] utilized Ti-6Al-4V ELI discs and subjected them to treatment with pulsed DC power for 3 min at 280 V in electrolytes containing bioactive ions such as Si, Zn, Mg, Mn, Sr, Ca, and P. The energy-dispersive X-ray spectroscopy (EDS) results revealed that the (Ca + Mg + Si + Mn + Zn + Sr)/P ratio was approximately 1.67. Some researchers have directly added HA to the electrolyte [100,101]. Zhou et al. [101] dispersed hydroxyapatite nanotubes (HNTs) in the MAO electrolyte and embedded HNTs into the Ti surface. When comparing three types of surfaces, Ti, MAO, and MAO-HNT, the HNT group outperformed the Ti and MAO groups in terms of angiogenesis, as evidenced by enhanced cell migration, tube formation, and vascular gene expression in HUVECs. Furthermore, concerning osteogenesis, the HNT group exhibited the highest levels of alkaline phosphatase (ALP) activity, collagen secretion, mineralized calcium nodules, and osteogenic gene expression in MC3T3-E1 cells, surpassing the other two groups. These findings suggest that HNT specimens can significantly promote angiogenesis and osteogenesis at both the cellular and molecular levels.

A good blood supply is essential for successful osseointegration. Studies have found that adulteration of appropriate amounts of Ca [101], Li [102], Se [103], Sr [104], Si [105], and Zn [106] elements can contribute to angiogenesis. Yu et al. [106] introduced zinc acetate into the electrolyte to produce coatings enriched with zinc. Co-cultivation of HUVECs with these coatings demonstrated excellent biocompatibility. This study explores the inherent connection between angiogenesis and osteogenesis, presenting evidence that Zn²⁺ promotes both processes.

MAO is a versatile technique that enhances the hardness, corrosion resistance, wear resistance, and biocompatibility of metal surfaces. This versatility makes it ideal for integration with other technologies to fine-tune surface properties for diverse application needs. The joint application of MAO and other surface modification technologies can incorporate functional elements into the MAO coating, such as hydrothermal treatment (HT) [107,108], dip coating [109,110], radiofrequency magnetron sputtering (RF-MS) [111], etc. Many research groups are exploring post-processing techniques using inorganic compounds to mimic the Ca/P ratio in a natural bone structure for enhanced bioactivity [50].

For instance, Song et al. [107] applied HT in two different solutions after MAO treatment, resulting in the formation of HA with a higher Ca/P ratio than the MAO-treated sample, facilitating the migration of Ca²⁺ and PO₄³⁻ ions to the surface. MAO and HT treatments provide binding sites for other functional factors like osteogenic growth factors, and antibacterial and anti-inflammatory substances [112]. In a study [113], MAO was initially used to create a hierarchical micro/nano-topography on a substrate, followed by electrochemical reduction in an alkaline solution. The results demonstrated improved BMSC adhesion, proliferation, and up-regulation of osteogenesis-related gene expression, indicating the potential to enhance osteogenic processes.

Magnetron sputtering, when combined with MAO, deposits metal particles onto coatings, enhancing osteogenesis. Park et al. [114] used RF-MS to create a Mn coating on the MAO-treated Ti-29Nb-xHf alloy, highlighting the potential for introducing metal coatings to boost surface osteogenic properties.

The combination of multiple surface modification techniques offers a broader selection of surface ingredients, including growth factors, hormones, and proteins. Bone morphogenic proteins (BMPs), such as BMP-2 and BMP-7, are pivotal for bone cell proliferation, differentiation, and matrix synthesis. Fibroblast growth factors (FGFs) like FGF-2 and FGF-7 influence bone cell proliferation and differentiation, while insulin-like growth factors (IGFs) like IGF-I and IGF-II promote cell growth and protein synthesis, crucial for bone tissue growth and repair. Vascular Endothelial Growth Factor (VEGF), as a key growth factor, directly stimulates new blood vessel formation. These biological factors are typically incorporated onto material surfaces through gentle methods, such as the dip-coating method [115,116], immersion [110], and chemical deposition [109], to ensure the preservation of their activity.

Teng [117] employed a method combining MAO and Ca, P layers with BMP co-precipitation to treat 3D-printed porous Ti alloy implants, named MAO-CaP-BMP2. The porosity facilitates bone tissue and blood vessel ingrowth. The microporous dioxide layer, created by MAO treatment, serves as nucleation sites for concurrent deposition of the CaP layer and BMP-2. This microstructural arrangement allows the extended, controlled release of BMP-2 over 35 days. The modification of Ti alloy implants with MAO-CaP-BMP2 enhances their osteoinductive and osteoconductive properties, leading to superior osteogenic and angiogenic outcomes.

3.2. Antibacterial

Among common orthopedic infections, osteomyelitis stands out, often stemming from pathogens like Staphylococcus aureus (S. aureus), Escherichia coli (E. coli), and Pseudomonas aeruginosa (P. aeruginosa). Symptoms encompass a range of discomforts such as pain, swelling, pus formation, the presence of fistulas and/or sinus tracts, and wound reopening. The approach to treatment depends on various factors including the severity of infection, the responsible pathogen, and the overall health of patients. Treatment varies based on infection severity, pathogen, and patient health, typically involving debridement to remove infected bone tissue and implanted devices. Debridement surgery is complex and requires an extended recovery period. Dental-implant-associated microbial colonization may give rise to peri-implant mucositis or even peri-implantitis. Key pathogens involved include Porphyromonas gingivalis (P. gingivalis), Streptococcus sanguinis (S. sanguinis), Aggregatibacter actinomycetemcomitans (A. actinomycetemcomitans), etc. Peri-implantitis has a lower incidence compared to peri-implant issues in orthopedic implants, but it still demands attention as it could directly contribute to implant failure [118].

Microbial cells adhering to surfaces can come together to form highly structured microbial communities commonly known as “biofilms”. These biofilms often display drug resistance, complicating infection treatment. The process of biofilm formation involves various stages, including the adsorption of proteins, microbial adhesion, aggregation, biofilm maturation, and the diffusion of biofilm (Figure 6). Proteins adhere to surfaces, facilitating microbial attachment through electrostatic interactions and protein receptor recognition. As attachment progresses, biofilm formation occurs as microorganisms aggregate. Once established, biofilms involve intercellular communication and extracellular polymeric substance (EPS) secretion, transforming into a three-dimensional structure [119].

Antimicrobial implant materials must inhibit biofilm formation, particularly in the initial stages. Researchers and clinicians continue to seek effective prevention and treatment of implant-associated infections. In recent years, the development of surface antibacterial methods using MAO has been a research hotspot in preventing and managing biofilm growth on Ti-based implants [120].

Some researchers suggest that MAO-generated anatase TiO₂ structures can reduce bacterial adhesion or confer antibacterial properties. For instance, Alipal et al. [121] found that circular TiO₂ structures on MAO coatings exhibit improved biocompatibility and antibacterial properties compared to needle-like TiO₂. Similarly, Jia et al. [122] proposed that the structure of TiO₂ layers created by MAO can trap and kill bacteria on implant surfaces. However, conflicting views exist in the literature, with some researchers arguing that rough, porous MAO coatings increase microbial adhesion [123,124], while others contend there is no significant difference compared to untreated Ti [125]. Current research on MAO-based antibacterial coatings mainly focuses on incorporating antibacterial additives to alter surface chemistry and enhance functionality.

The main mechanisms to promote the killing of bacteria on the surface of antimicrobial implants are (i) the contact reaction of bacteria with the surface, causing the destruction of bacterial membranes or inhibiting the synthesis of their membranes; (ii) the gradual release of ions from the coating, resulting in electron transfer and cell membrane destruction; (iii) antimicrobial ions enter microbial cells, interrupt protein synthesis, block Adenosine Triphosphate (ATP) synthesis and deoxyribonucleic acid (DNA) replication, and promote apoptosis; and (iv) generation of ROS, which also induces apoptosis by denaturing proteins and damaging DNA [126]. Costa et al. [119] categorized antimicrobial agents into four primary groups: organic (biopolymers and bioactive agents), inorganic (metallic and semi-metallic elements), drugs (substances with pharmacological action), and biological (growth factors, peptides, and extracellular matrix proteins). Among these additives, people are more interested in inorganic elements, due to their stability and cost-effectiveness. The specific effects of elements on the biological function of the coating are shown in Table 3.

Since 2009, one-step MAO has been employed to create antimicrobial Ti surfaces by binding Ag, Cu, and Zn elements, as summarized in 35 articles by Shimabukuro up to 2020 [51]. The half-maximal inhibitory concentrations (IC50) of Ag⁺, Cu²⁺, and Zn²⁺ on MC3T3-E1 cells are 2.77, 15.9, and 90.0 μM, respectively [127]. Ag exhibits high toxicity, but controlled dosages do not negatively affect osteoblast activity. Researchers are increasingly interested in incorporating diverse elements (Ag + Cu [23], Ag + Zn [128,129,130,131], Cu + Zn [132,133]) into coatings through one- or two-step MAO processes to enhance antimicrobial effects or/and extend the duration of antimicrobial activity. Shimabukuro et al. [134,135] simulated the biodegradation behavior of Ti doped with Ag, Cu, and Zn by MAO soaked in physiological saline for 28 days to study the changes in surface composition and antibacterial effect. After 28 days, the antibacterial effect of the Ag-incorporated samples weakened, the antibacterial effect of Cu was maintained, and the antibacterial effect of Zn enhanced in the later period (Figure 7). In another study, Tsutsumi [136] employed a two-step MAO process to tailor surface coatings with Ag, with or without Zn, to assess the binding and ion release behavior of Ag and Zn in the resulting oxide layer. Over a 6-month immersion period, the release of Ag⁺ from the oxide layer indicated early antibacterial performance, while the release of Zn²⁺ signified later antibacterial effectiveness.

Research involving Mn on the surface of Ti-based implants shows promise, as it enhances osteoblast activity while also inhibiting the proliferation of Gram-negative E. coli and P. aeruginosa. Zhao et al. [137] employed MAO to create a Mn-TiO₂ microporous coating. Their results revealed that the coating induced E. coli cell wall perforation, effectively inhibiting bacterial proliferation. Additionally, Bi compounds have shown efficacy in treating mucosal and dermal infections and are recognized as antimicrobial additives for calcium phosphate bone cement. Lin et al. [138] used MAO to develop a TiO₂ coating containing the bismuth (Bi) element. The coating with 6.2% Bi displayed good biocompatibility with MG63 cells and demonstrated 1.5 times higher antibacterial efficacy against A. actinomycetemcomitans and 1.9 times higher efficacy against methicillin-resistant Staphylococcus aureus (MRSA) compared to the control group. MAO technology can also incorporate other non-metallic antibacterial ingredients into the coating, such as Ce [139], Se [103], I [140,141], B [142,143], and F [144,145,146].

MAO combined with other techniques can produce antimicrobial coatings that enhance effectiveness and minimize cytotoxicity. Arun et al. [111] employed a combination of MAO and RF-MS processes to create dual-phase coatings on Ti-6Al-4V. They conducted MAO in Ca/P electrolytes with varying Ag NP concentrations and applied an HA top layer using the RF-MS process. This approach generated a passivation layer with robust barrier properties. Ag contributed to antibacterial activity, while HA promoted cell proliferation and mitigated the minor adverse effect of Ag on cell viability. To enhance cell compatibility with antibacterial coatings or extend antibacterial substance release, polydopamine [147,148] and polylactic acid [149] can be loaded on the coating surface through secondary surface treatment.

Graphene and its derivatives (e.g., graphene oxide (GO) and redox graphene (rGO)) can be integrated into MAO coatings either directly through electrolytes or in combination with techniques like electrophoretic deposition (EPD), sol-gel, and dropping coating. This incorporation imparts favorable antibacterial and tribological properties [150]. For instance, Mazinani et al. [151] combined MAO and EPD to deposit GO sheets onto the MAO coating, resulting in significant antimicrobial properties against E. coli and S. aureus. Moreover, San et al. [152] created a suspension of rGO and Ag NPs, added it to the MAO electrolyte, and achieved a more effective coating against MRSA due to the generation of ROS and the “Nanoknife” structure of rGO.

Chitosan (CS) interacts with the anionic components on the microbial cell wall, resulting in cell wall damage and the killing of microorganisms, due to its cationic nature and polyamine structure. CS can be fixed on the MAO coating by dip coating [153]. For example, BMP-2-encapsulated CS gives the coating good osteogenic and antibacterial properties [115].

Incorporating drugs, such as antibiotics, is an effective means of rapid sterilization during early implantation stages. For example, Zhou et al. [154] used ciprofloxacin-loaded CS hydrogel to modify the coating, effectively sterilizing the sample with the bacteriostatic zone. Xu et al. [155] covalently grafted vancomycin to the MAO coating, creating Ti-@MV with significant antibacterial properties and enhanced osteogenic differentiation of BMSCs. In another study [156], octenidine (OCT) was loaded onto MAO-prepared coatings via electrophoresis, resulting in improved early antibacterial efficiency without compromising cell compatibility. This early-stage (4–6 h) intervention is crucial for preventing long-term bacterial infections. Octenidine, vancomycin, and heparin were among the early attempts at drug loading on polyethylene glycol MAO-treated surfaces. However, their initial release is too large, and it is easy for bacteria to develop drug resistance. Therefore, the use of antibiotics in implant surface modification is gradually decreasing [157]. In terms of biological compounds, antimicrobial peptides (AMPs) have become the main substances for functionalization in Ti surfaces through a multi-step process on MAO surfaces [147,158]. However, the main challenge with these antimicrobial surfaces is that cross-linking of peptides on the surfaces is difficult, resulting in less stable coatings, limited duration of antimicrobial activity, and higher manufacturing costs [159].

Phototherapy, including photothermal and photodynamic therapy, offers safe and controllable methods. Photothermal therapy employs agents that generate high temperatures under near-infrared light, effectively disrupting bacterial structures. In contrast, photodynamic therapy relies on photosensitizers absorbing specific light wavelengths to create reactive oxygen species (ROS), which disrupt bacterial DNA, proteins, and biomembranes. Hydrothermal treatment or sol-gel can be used to combine photothermal agents or photosensitizers into the coating [160]. Chai et al. [161] successfully prepared a TiO₂/MoO₂/chitosan coating on Ti surfaces using a composite process involving MAO, HT, and electrospinning. The MoO₂ significantly improved the phototherapy performance. Both in vitro and in vivo experiments demonstrated excellent antibacterial capabilities against Streptococcus mutans (S. mutans) after 15 min of irradiation with near −808 nm wavelength infrared light, facilitated by the synergistic effects of a high temperature and ROS. Moreover, Li et al. [108] used MAO and HT techniques to form a β- FeOOH/Fe TiO₂ heterojunction, which can also generate ROS under irradiation. In addition, ultraviolet rayscan remove surface carbon contaminants and increase the negative charge density on the surface of crystal TiO₂, thereby reducing bacteria through electrostatic repulsion [125]. However, because UV rays have carcinogenic effects, they can be used as a pretreatment before implant placement.

3.3. Anti-Inflammatory

The biological properties of materials are initially evaluated through in vitro studies on osteoblast activity. In some cases, differences between in vivo and in vitro research results may stem from immune cell responses. Bone immunology emphasizes the role of immune cells in osteointegration, which involves the inflammatory response in the early and late stages after implantation [162].

The implantation of biomaterials is intricately tied to a continuous biological process (Figure 8). Shortly after implantation, proteins from the vascular system and interstitial fluid adhere to the surface of the biomaterial within minutes. Subsequently, various cells, including platelets, monocytes, and macrophages, infiltrate and adhere to the biomaterial. Within 1 to 5 days post-implantation, immune cells release cytokines on the biomaterial surface, initiating a pro-inflammatory response while recruiting bone repair cells (e.g., BMSCs and OBs) from distant sites to the inflammatory site [163]. This early-stage inflammation is crucial. The fate of the bone biomaterial, whether it becomes encapsulated by a fibrous capsule (indicating failure) or undergoes long-term replacement by new bone (the desired outcome), is determined by bone repair cell interactions with the inflammatory microenvironment and biomaterial properties in the following days. Acute inflammation typically subsides in less than 1 week, while a duration exceeding 3 weeks may lead to implant failure. The emerging field of bone immunology and immune modulation has shifted our expectations for bone biomaterials, emphasizing their role in bone immune modulation rather than just reducing inflammatory responses [164].

Recent advancements in Ti surface engineering focus on managing host and inflammatory responses. This approach leverages the benefits of physiological inflammation, such as cellular debris removal and infection prevention, while also mitigating the risks associated with severe inflammatory reactions. These risks include implant rejection, early or late failures, and impacts on implant functionality.

To modulate the inflammatory response of biomaterials, these strategies can be followed [164]:

• Adsorption of specific proteins: control the adsorption of plasma proteins on the implant surface to influence inflammatory responses.
• Macrophage polarization: M1 (pro-inflammatory) and M2 (anti-inflammatory) phenotypes can be induced. M1 macrophages clear debris and potential pathogens in the early stages of implantation, while M2 macrophages support inflammation resolution and tissue healing. The M1 phenotype is produced by M0 macrophages activated by the lipopolysaccharide (LPS) or interferon-γ (INF-γ). It secretes inflammatory mediators, including Interleukin-1β (IL-1β), IL-6, IL-12, IL-23, ROS, and Tumor Necrosis Factor-α (TNF-α). However, the M2 phenotype is induced by stimulation with IL-4, IL-10, and IL-1β. M2 macrophages release anti-inflammatory cytokines and growth factors, such as IL-10 and transforming growth factor-β (TGF-β).
• Surface topography: Surface roughness and texture can affect macrophage phenotype. Rough surfaces may favor M1 activation, while nano- or micron-scale directional textures promote an M2-like response. Hydrophilic surfaces can reduce pro-inflammatory cytokines.
• Biomimetic coating: modify the implant surface with biomimetic coatings like collagen, transmembrane molecular markers, mesenchymal stem cells, or biological peptides to reduce macrophage activation and pro-inflammatory factors.
• Local drug delivery: implement local drug delivery systems on the implant surface to create a localized anti-inflammatory environment with high drug concentration and minimal systemic side effects.

Local inflammatory reactions can result from debris accumulation caused by wear during implantation or long-term friction between the implant and surrounding bone tissue [165,166]. Furthermore, the corrosion of metals in the complex environment of the body can release metal ions, potentially triggering inflammation. MAO coatings offer excellent tribological properties and corrosion resistance. They can be modified to regulate host and inflammatory responses by adjusting the surface topography and composition of the implant. Li et al. [163] examined the in vitro immune response to Ca- and Si-doped MAO coatings. They found that the macrophage interaction with MAO-Ti favored the M1 phenotype, but collagen synthesis and matrix mineralization for osteocyte-like cell growth on the coating were favorable. This suggests that MAO-modified implant–bone interfaces support osteogenesis both before and after osteoblast recruitment to the biomaterial surface. Furthermore, Ti-Ta metal–metal composites with bioceramic coatings containing Ca and Si also demonstrate osteogenic, anti-inflammatory, and bone immunoregulatory properties [167].

Some metallic elements, such as Cu [168], Zn [169], Mg [170], Li [102], and Co [171], have been shown to modulate the immune environment at the implant surface. They can regulate macrophage polarization phenotypes, thus promoting either early pro-inflammatory and antimicrobial effects or long-term anti-inflammatory and pro-osteogenic effects of the implant. For instance, Li et al. [168] utilized MAO to create Cu-containing ceramic coatings on Ti surfaces (Cu-MAO). Macrophages on Cu-MAO displayed polarization toward the M1 phenotype, exhibiting high levels of inducible nitric oxide synthase (iNOS) expression, low arginase-1 (Arg1) expression, enhanced release of IL-6, and inhibited release of IL-4 and IL-10. This promoted a favorable inflammatory microenvironment for osteoblast-like cell differentiation and enhanced bactericidal abilities of the surface. Sun et al. [169] took a similar approach by developing Zn-doped MAO coatings on Ti surfaces, which exhibited excellent biocompatibility and osteogenic differentiation performance with BMSCs. In addition, evaluation of RAW264.7 cells’ survival on the surface and related pro-inflammatory markers revealed that this coating downregulated the expression of pro-inflammatory genes. Peng et al. [102] incorporated Lithium (Li) into Ti surface coatings via MAO. Low-dose Li was slowly released as ions, enhancing the recruitment of BMDMs while favoring M2 polarization and limiting M1 polarization. In vivo, implantation mimicking aseptic loosening revealed the ability of Li to mitigate inflammatory responses. Additionally, Li-doped MAO coatings supported healthy growth of both mouse embryonic cell lines (C3H10T1/2) and HUVECs in a macrophage-conditioned medium, highlighting their potential for osteogenic differentiation and angiogenesis. The molecular mechanism involved the regulation of cascade molecules in the PI3K/AKT signaling pathway for bone immunomodulation. Similarly, Yang et al. [171] prepared Co-doped MAO coatings on Ti implants with varying Co content. Their results showed that, in comparison to pure Ti samples, cobalt-loaded Ti exhibited immunomodulatory functions on macrophages, upregulating the expression of M1 genes and downregulating the expression of M2 genes. Notably, higher cobalt levels induced the polarization of macrophages into the M2 type.

Similar to antimicrobial properties, other surface modification techniques can incorporate a greater amount of anti-inflammatory components onto the surface of the MAO coating. Bai et al. [172] investigated the effects of MAO and steam heat treatment (SHT) on the preparation of nano-HA coatings on porous Ti surfaces. They found that Ti surfaces coated with HA NPs exhibited adjustable inflammatory responses, while HA nanorods negatively affected osteo-/angiogenesis and osteoimmunomodulation. Additionally, they [100] annealed the MAO-HA surfaces and observed that annealing temperature influenced surface topography, wettability, and chemical properties. In vitro experiments demonstrated that MAO coatings annealed at 650 °C promoted the proliferation and differentiation of osteoblasts and endothelial cells, while inhibiting the inflammatory response of macrophages.

Schmidlin et al. [173] found that ZrO₂ particles induced lower expression of inflammatory factors compared to TiO₂ particles at an equivalent dose. Zhu et al. [174] combined the MAO and sol-gel methods to fabricate a multifunctional composite coating on the surface of Ti, composed of SiO₂ particles and zirconium hydrogen phosphate (ZrP). The ZrP coating exhibited superhydrophilicity and significantly reduced the release of pro-inflammatory cytokines (TNF-α, IL-6, and IL-1β). This coating also enhanced corrosion resistance and friction, resulting in reduced Ti ion/fragment release and minimized inflammatory responses. Wang et al. [175] prepared Sr-doped nanostructures on the Ti surface through MAO and electrochemical deposition, and subsequently incorporated silk fibroin protein-based nanocomposites (Ti-MAO/Sr/LBL/WNP) through a layer-by-layer self-assembly technique (LBL). This coating exhibited sustained release of wogonin and Sr ions for over 7 days, promoted the expression of anti-inflammatory factors (TGF-β1 and Arg-1) in M2 macrophages, and facilitated cell proliferation and osteogenic differentiation. Moreover, MAO coatings can be utilized to locally load bioactive substances and drugs, such as peptides that recruit immune-modulating cells (T cells, monocytes, mast cells, and neutrophils) [158].

MAO provides biofunctional coatings that interact with tissues over an extended period, involving multiple biological and biochemical processes [90]. This interaction is determined by the implant characteristics and the implantation environment. Evaluating the biocompatibility and long-term performance of the coatings is a task that requires various factors and methods. Biocompatibility testing includes both in vitro and in vivo assessments [109,113]. In vitro experiments can detect the effects of the coating on cell adhesion, proliferation, and differentiation. In vivo experiments assess the biocompatibility, tissue response, and implant stability within live animals. Assessing engineering parameters in the surrounding tissues, such as vascularization, bone formation, and healing, provides information on the coating impact on tissue reconstruction. Monitoring implant position and stability through imaging techniques such as X-rays, CT scans, and MRI is crucial for detecting issues like coating delamination, implant loosening, or displacement. Animal experiments observe the morphology, structure, chemical composition, and mechanical properties of the coating after a certain period of implantation, offering insights into the long-term stability of coatings in vivo. Evaluating long-term performance also involves considering inflammation and immune reactions around the implant, monitoring through tissue slices, inflammation levels, biomarker detection, and immune cell activity [168,172,175,176,177] (refer to the literature listed in Table 3). Integrating these assessments provides a more comprehensive understanding of the performance of MAO coatings during long-term use and their interaction with surrounding tissues. This multidisciplinary task typically requires collaboration among professionals in medicine, biomedical engineering, and materials science. Moreover, dental implants and hip stems featuring MAO coatings have been introduced as products and subjected to market testing over the long term, yet relevant literature reports are scarce [178,179].

Table 3. Coatings contain different components with different biological functions and are validated in vitro and in vivo.

Number	Biologic Function	Adding Substances	Technique	In Vitro		In Vivo	Conclusion/Remark	Reference
				Cell	Bacterium	Animal Parts
1	Osteogenesis	Ca, Sr	MAO	hBMSCs			Simultaneously incorporating Ca and Sr demonstrated superior promotion of hBMSC proliferation.	[97]
2	Osteogenesis/Angiogenesis	Zn	MAO	HUVECs and BMSCs			In the Zn2+ environment, angiogenesis and osteogenesis mutually promote each other.	[106]
3	Osteogenesis/Angiogenesis	Hydroxyapatite nanotubes (HNTs)	MAO	HUVECs and MC3T3-E1 cells			HNT specimens promote both angiogenesis and osteogenesis on cellular and molecular levels.	[101]
4	Osteogenesis	B	MAO, hydrothermal treatment, and heat treatment	SaOS-2 cells			Nanorods inhibit SaOS-2 cell activity, whereas nanoparticles promote it.	[143]
5	Osteogenesis	Hierarchical coatings	MAO, electrochemical reduction	BMSCs		Beagle dogs, the shaft of the canine femur	The hierarchical coatings show higher osteogenesis rates compared to the ordinary MAO group.	[113]
6	Osteogenesis	HA, BMP-2	MAO, dip coating	MC3T3-E1 cells		Beagle femur	The interface bonding strength between HA/BMP-2 coating and surrounding new bone tissue is higher than that of Ca/PMAO coating.	[109]
7	Osteogenesis/Angiogenesis	Ca, P, BMP-2	3D printing, sandblasting etching, MAO, electrochemical deposition	BMSCs		New Zealand White Rabbit Skull	MAO-CaP-BMP-2 is superior to the MAO and MAO-CaP groups in new bone formation.	[117]
8	Osteogenesis/Antibacterial	Ca, P	MAO	hFOBs	E. coli and S. aureus		Volcanic-crater-like and needle-like CaP structures form at 350 V and 450 V, respectively. The former exhibits superior antibacterial performance and biocompatibility.	[121]
9	Bioactivity/Antibacterial	Ca, P	MAO, UV catalysis	HGFs	S. sanguinis		Photofunctionalization reduces hydrocarbons and enhances surface protein adsorption.	[125]
10	Osteogenesis/Antibacterial	Zn	MAO	MC3T3-E1 cells	E. coli		Incubation with salt solution converts Zn ions into zinc oxide, which helps with long-lasting antibacterial activity.	[134]
11	Antibacterial/Osteogenesis/Angiogenesis	Sr, Zn	MAO	HUVECs, BMSC	MRSA and P. gingivalis	Rat femoral model	The surface osteogenesis of samples doped with Sr and Zn is superior to other groups. (No in vivo antibacterial test conducted.)	[104]
12	Antibacterial	Ag, Cu NPs	MAO	MC3T3-E1 cells	MRSA	Mouse femur ex vivo experiment	Ag and Cu ions synergistically kill bacteria, allowing a 10-fold reduction in Ag ion concentration with consistent antibacterial efficacy.	[124]
13	Osteogenesis/Antibacterial	Ag, Zn	3D printing, MAO	MC3T3-E1 cells	MRSA	Mouse femur ex vivo experiment	The synergistic effect of Ag and Zn reduces the concentration of Ag+ by 120 times.	[128]
14	Osteogenesis/Antibacterial	Ag, Zn	MAO	MC3T3-E1 cells	S. aureus		Ag and ZnO synergy enhances antibacterial performance and promotes CaP phase formation.	[129]
15	Osteogenesis/Antibacterial	Ag, Zn	MAO	MC3T3-E1 cells	S. aureus		Ag and Zn ion release is above the antibacterial threshold yet well below cytotoxic levels.	[130]
16	Osteogenesis/Antibacterial	Ag, Zn	MAO		S. aureus		Ag and Zn have good synergistic antibacterial effects.	[131]
17	Osteogenesis, Antibacterial	Cu, Zn	MAO	MG63	E. coli, S. aureus, and MRSA		Orthogonal experiments explore electrolyte effects on coatings, with phytic acid supplying the P element.	[132]
18	Skin-integration/Antibacterial	Cu, Zn	MAO	Fibroblasts (L-929)	S. aureus		The synergistic effect of Cu and Zn facilitates skin integration and antibacterial activity.	[133]
19	Osteogenesis, Anti-tumor/Antibacterial	Se	MAO	BMSCs, cancerous osteoblasts	S. aureus and E. coli		Se doping enhances osteogenic, anti-tumor, and antibacterial properties.	[103]
20	Osteogenesis/Antibacterial	Mn	MAO	MC3T3-E1 cells	E. coli	Rabbit femur	The coating induces osteogenesis and promotes osseointegration.	[137]
21	Antibacterial	Bi	MAO	MG63 cells	A. actinomycetemcomitans, MRSA		Bismuth nitrate has excellent antibacterial activity compared to bismuth acetate, bismuth gallate, and silver nitrate.	[138]
22	Osteogenesis/Antibacterial	Ce	MAO	BMSCs	P. gingivalis, S. aureus	Osteoporotic rat hind legs	Ce-TiO2 coating has excellent antibacterial and anti-inflammatory properties.	[139]
23	Antibacterial	I	MAO, HT, photocatalysis	BMSCs	S. aureus	Tibial Intramedullary Infection Model of Rats	Under NIR, the coating has good antibacterial and osteogenic properties.	[140]
24	Antibacterial	I	MAO, electrophoresis	BMSCs	S. aureus and E. coli	The rat osteomyelitis intramedullary nail model	Thirty days after implantation, excellent antimicrobial ability was verified.	[141]
25	Bioactivity/Antibacterial	B	MAO	ADSCs	S. aureus and P. aeruginosa		Add a small amount of sodium tetraborate to the Ca, P electrolyte system.	[142]
26	Osteogenesis/Antibacterial	F	MAO	BMSCs	S. aureus and E. coli	Rabbit femur	Coatings with high F addition showed improved antibacterial and osteogenic abilities.	[144]
27	Antibacterial/Osteogenesis/Angiogenesis	Sr, Co, and F	MAO	BMSCs	S. aureus and E. coli	Rabbit femur	Sr, Co, and F co-doped coatings induce osteogenesis.	[145]
28	Osteogenesis/Antibacterial	Mn, F	MAO	BMSCs	S. aureus		Mn and F co-doped coatings show excellent wear and corrosion resistance, along with strong antibacterial properties.	[146]
29	Osteogenesis/Antibacterial	Cu, BMP-2	MAO, dip coating	MC3T3-E1 cells	E. coli, MRSA, Neurospora crassa, and Candida albicans	Mouse craniotomy model	The coating significantly promotes osseointegration.	[110]
30	Osteogenesis/Antibacterial	Ag, HA	MAO, RF-MS	MC3T3-E1 cells	E. coli		This coating exhibits strong biological activity and antibacterial properties.	[111]
31	Bioactivity/Antibacterial	Ag NPs, polylactic acid (PLA)	MAO, electrospinning	MC3T3-E1 cells	S. aureus		PLA ultrafine fibers produced by electrospinning can control the release of silver ions.	[149]
32	Osteogenesis/Antibacterial	AgNPs, polydopamine	MAO, dip coating	MG63 cells	S. aureus	New Zealand rabbit subdermal implantation	This coating exhibits strong biological activity and antibacterial properties.	[122]
33	Osteogenesis/Antibacterial	Polydopamine, cationic antimicrobial peptide LL-3, phospholipid	MAO, dip coating	BMSCs and OBs	S. aureus and E. coli		The coating exhibits good osteogenesis and antibacterial properties.	[147]
34	Antibacterial	GO	MAO, EPD		S. aureus and E. coli		Achieves ~80% antibacterial activity against E. coli and 100% against S. aureus.	[151]
35	Antibacterial	rGO, Ag NPs	MAO	MC3T3-E1 cells	MRSA		The coating exhibits good osseogenesis and antibacterial properties.	[152]
36	Osteogenesis/Antibacterial	HA, chitosan (CS)	MAO, dip coating	MC3T3-E1 cells	E. coli		Higher usage of CS results in decreased biological performance but improved antimicrobial performance.	[153]
37	Osteogenesis/Antibacterial	HA, CS hydrogel containing ciprofloxacin	MAO, HT, chemical grafting	hBMSCs	S. aureus and E. coli		The coating exhibits good osseogenesis and antibacterial properties.	[154]
38	Osteogenesis/Antibacterial	BMP-2/CS/HA	MAO, dip coating	MC3T3-E1 cells	E. coli		CS encapsulation sustains BMP-2 release with added antibacterial properties.	[115]
39	Antibacterial	Vancomycin	MAO, dip coating			The rabbit osteomyelitis model (infection with MRSA)	In vivo studies demonstrate the potential of this coating to prevent MRSA infection.	[148]
40	Osteogenesis/Antibacterial	Vancomycin	MAO, dip coating, chemical grafting	BMSCs	S. aureus	Rat femur	Functional coatings prevent prosthesis infection and promote bone integration at the interface.	[155]
41	Antibacterial	Mesoporous silica NPs (MSNs), octenidine (OCT)	Electrophoretic-enhanced MAO	OBs	S. aureus and E. coli		The coating exhibits good osseogenesis and antibacterial properties.	[156]
42	Bioactivity/Antibacterial	N, Bi	MAO, photocatalysis	HGFs	Streptococcus sanguinis and Actinomyces nasseri		The coating has bactericidal properties under visible light.	[123]
43	Osteogenesis/Antibacterial	MoSe2, CS	MAO, electrospinning, photocatalysis	MC3T3-E1 cells	S. mutans	Rat tibia	Adding MoSe2 significantly enhances TiO2 coating photothermal and photodynamic capabilities.	[161]
44	Skin-integration/Antibacterial	β-FeOOH, Fe-TiO2	MAO, HT, photocatalysis	Mouse fibroblasts (L-929)	S. aureus	Mouse skin infection model	The β-FeOOH/FeTiO2 heterojunction prevents bacterial infection under light irradiation.	[108]
45	Osteogenesis/Anti-inflammatory	Ca, Si	MAO	SaOS-2 cells			The coating inhibits inflammation and induces M2 macrophage polarization.	[167]
46	Antibacterial/Immunoregulation	Cu	MAO	RAW 264.7 macrophages, SaOS-2 cells	S. aureus		Cu boosts macrophage-driven osteogenesis and antibacterial activity in biomaterials.	[168]
47	Osteogenesis/Anti-inflammatory	Zn	MAO	RAW264.7 macrophages, BMSCs			The coating shows good osteogenic and anti-inflammatory properties.	[169]
48	Osteogenesis/Anti-inflammatory	Mg	MAO	RAW 264.7 macrophages			Mg acts as an anti-inflammatory agent, inhibiting inflammation and promoting osteogenesis.	[170]
49	Anti-inflammatory	Co	MAO	RAW 264.7 macrophages		Mouse air chamber model	Cobalt-loaded Ti exhibits immune-regulatory effects on macrophages.	[171]
50	Osteogenesis/Angiogenesis/Anti-inflammatory	Li	MAO	BMDMs, mouse embryonic cell line (C3H10T1/2), HUVEC		Mouse air-pouch model	Low Li doses effectively regulate immunity, and promote osteogenesis.	[102]
51	Osteogenesis/Angiogenesis/Anti-inflammatory	HA	MAO, SHT	MC3T3-E1 cells, human umbilical vein fusion cells, RAW 264.7 cells		Rabbit femur	This coating promotes osteogenesis and angiogenesis, and induces M2 macrophage phenotype.	[172]
52	Osteogenesis/Anti-inflammatory	HA	MAO, SHT	MC3T3-E1 cells, endothelial cells, RAW 264.7 cells		Rabbit femur	Nanoparticle-shaped HA is beneficial for osteogenesis, angiogenesis, and immune regulation, whereas nanorod-shaped HA is the opposite.	[173]
53	Osteogenesis/Anti-inflammatory	SiO2, ZnPs	MAO, sol-gel	MC3T3-E1 cells			The coating shows good osteogenic and anti-inflammatory properties.	[175]
54	Osteogenesis/Anti-inflammatory	Sr, silk fibroin-based wogonin NPs	MAO, electrochemical deposition, LBL	RAW 264.7 cells, OBs		Osteoporotic rat femur	The coating shows good osteogenic and anti-inflammatory properties.	[176]

Table 3. Coatings contain different components with different biological functions and are validated in vitro and in vivo.

Number	Biologic Function	Adding Substances	Technique	In Vitro		In Vivo	Conclusion/Remark	Reference
				Cell	Bacterium	Animal Parts
1	Osteogenesis	Ca, Sr	MAO	hBMSCs			Simultaneously incorporating Ca and Sr demonstrated superior promotion of hBMSC proliferation.	[97]
2	Osteogenesis/Angiogenesis	Zn	MAO	HUVECs and BMSCs			In the Zn2+ environment, angiogenesis and osteogenesis mutually promote each other.	[106]
3	Osteogenesis/Angiogenesis	Hydroxyapatite nanotubes (HNTs)	MAO	HUVECs and MC3T3-E1 cells			HNT specimens promote both angiogenesis and osteogenesis on cellular and molecular levels.	[101]
4	Osteogenesis	B	MAO, hydrothermal treatment, and heat treatment	SaOS-2 cells			Nanorods inhibit SaOS-2 cell activity, whereas nanoparticles promote it.	[143]
5	Osteogenesis	Hierarchical coatings	MAO, electrochemical reduction	BMSCs		Beagle dogs, the shaft of the canine femur	The hierarchical coatings show higher osteogenesis rates compared to the ordinary MAO group.	[113]
6	Osteogenesis	HA, BMP-2	MAO, dip coating	MC3T3-E1 cells		Beagle femur	The interface bonding strength between HA/BMP-2 coating and surrounding new bone tissue is higher than that of Ca/PMAO coating.	[109]
7	Osteogenesis/Angiogenesis	Ca, P, BMP-2	3D printing, sandblasting etching, MAO, electrochemical deposition	BMSCs		New Zealand White Rabbit Skull	MAO-CaP-BMP-2 is superior to the MAO and MAO-CaP groups in new bone formation.	[117]
8	Osteogenesis/Antibacterial	Ca, P	MAO	hFOBs	E. coli and S. aureus		Volcanic-crater-like and needle-like CaP structures form at 350 V and 450 V, respectively. The former exhibits superior antibacterial performance and biocompatibility.	[121]
9	Bioactivity/Antibacterial	Ca, P	MAO, UV catalysis	HGFs	S. sanguinis		Photofunctionalization reduces hydrocarbons and enhances surface protein adsorption.	[125]
10	Osteogenesis/Antibacterial	Zn	MAO	MC3T3-E1 cells	E. coli		Incubation with salt solution converts Zn ions into zinc oxide, which helps with long-lasting antibacterial activity.	[134]
11	Antibacterial/Osteogenesis/Angiogenesis	Sr, Zn	MAO	HUVECs, BMSC	MRSA and P. gingivalis	Rat femoral model	The surface osteogenesis of samples doped with Sr and Zn is superior to other groups. (No in vivo antibacterial test conducted.)	[104]
12	Antibacterial	Ag, Cu NPs	MAO	MC3T3-E1 cells	MRSA	Mouse femur ex vivo experiment	Ag and Cu ions synergistically kill bacteria, allowing a 10-fold reduction in Ag ion concentration with consistent antibacterial efficacy.	[124]
13	Osteogenesis/Antibacterial	Ag, Zn	3D printing, MAO	MC3T3-E1 cells	MRSA	Mouse femur ex vivo experiment	The synergistic effect of Ag and Zn reduces the concentration of Ag+ by 120 times.	[128]
14	Osteogenesis/Antibacterial	Ag, Zn	MAO	MC3T3-E1 cells	S. aureus		Ag and ZnO synergy enhances antibacterial performance and promotes CaP phase formation.	[129]
15	Osteogenesis/Antibacterial	Ag, Zn	MAO	MC3T3-E1 cells	S. aureus		Ag and Zn ion release is above the antibacterial threshold yet well below cytotoxic levels.	[130]
16	Osteogenesis/Antibacterial	Ag, Zn	MAO		S. aureus		Ag and Zn have good synergistic antibacterial effects.	[131]
17	Osteogenesis, Antibacterial	Cu, Zn	MAO	MG63	E. coli, S. aureus, and MRSA		Orthogonal experiments explore electrolyte effects on coatings, with phytic acid supplying the P element.	[132]
18	Skin-integration/Antibacterial	Cu, Zn	MAO	Fibroblasts (L-929)	S. aureus		The synergistic effect of Cu and Zn facilitates skin integration and antibacterial activity.	[133]
19	Osteogenesis, Anti-tumor/Antibacterial	Se	MAO	BMSCs, cancerous osteoblasts	S. aureus and E. coli		Se doping enhances osteogenic, anti-tumor, and antibacterial properties.	[103]
20	Osteogenesis/Antibacterial	Mn	MAO	MC3T3-E1 cells	E. coli	Rabbit femur	The coating induces osteogenesis and promotes osseointegration.	[137]
21	Antibacterial	Bi	MAO	MG63 cells	A. actinomycetemcomitans, MRSA		Bismuth nitrate has excellent antibacterial activity compared to bismuth acetate, bismuth gallate, and silver nitrate.	[138]
22	Osteogenesis/Antibacterial	Ce	MAO	BMSCs	P. gingivalis, S. aureus	Osteoporotic rat hind legs	Ce-TiO2 coating has excellent antibacterial and anti-inflammatory properties.	[139]
23	Antibacterial	I	MAO, HT, photocatalysis	BMSCs	S. aureus	Tibial Intramedullary Infection Model of Rats	Under NIR, the coating has good antibacterial and osteogenic properties.	[140]
24	Antibacterial	I	MAO, electrophoresis	BMSCs	S. aureus and E. coli	The rat osteomyelitis intramedullary nail model	Thirty days after implantation, excellent antimicrobial ability was verified.	[141]
25	Bioactivity/Antibacterial	B	MAO	ADSCs	S. aureus and P. aeruginosa		Add a small amount of sodium tetraborate to the Ca, P electrolyte system.	[142]
26	Osteogenesis/Antibacterial	F	MAO	BMSCs	S. aureus and E. coli	Rabbit femur	Coatings with high F addition showed improved antibacterial and osteogenic abilities.	[144]
27	Antibacterial/Osteogenesis/Angiogenesis	Sr, Co, and F	MAO	BMSCs	S. aureus and E. coli	Rabbit femur	Sr, Co, and F co-doped coatings induce osteogenesis.	[145]
28	Osteogenesis/Antibacterial	Mn, F	MAO	BMSCs	S. aureus		Mn and F co-doped coatings show excellent wear and corrosion resistance, along with strong antibacterial properties.	[146]
29	Osteogenesis/Antibacterial	Cu, BMP-2	MAO, dip coating	MC3T3-E1 cells	E. coli, MRSA, Neurospora crassa, and Candida albicans	Mouse craniotomy model	The coating significantly promotes osseointegration.	[110]
30	Osteogenesis/Antibacterial	Ag, HA	MAO, RF-MS	MC3T3-E1 cells	E. coli		This coating exhibits strong biological activity and antibacterial properties.	[111]
31	Bioactivity/Antibacterial	Ag NPs, polylactic acid (PLA)	MAO, electrospinning	MC3T3-E1 cells	S. aureus		PLA ultrafine fibers produced by electrospinning can control the release of silver ions.	[149]
32	Osteogenesis/Antibacterial	AgNPs, polydopamine	MAO, dip coating	MG63 cells	S. aureus	New Zealand rabbit subdermal implantation	This coating exhibits strong biological activity and antibacterial properties.	[122]
33	Osteogenesis/Antibacterial	Polydopamine, cationic antimicrobial peptide LL-3, phospholipid	MAO, dip coating	BMSCs and OBs	S. aureus and E. coli		The coating exhibits good osteogenesis and antibacterial properties.	[147]
34	Antibacterial	GO	MAO, EPD		S. aureus and E. coli		Achieves ~80% antibacterial activity against E. coli and 100% against S. aureus.	[151]
35	Antibacterial	rGO, Ag NPs	MAO	MC3T3-E1 cells	MRSA		The coating exhibits good osseogenesis and antibacterial properties.	[152]
36	Osteogenesis/Antibacterial	HA, chitosan (CS)	MAO, dip coating	MC3T3-E1 cells	E. coli		Higher usage of CS results in decreased biological performance but improved antimicrobial performance.	[153]
37	Osteogenesis/Antibacterial	HA, CS hydrogel containing ciprofloxacin	MAO, HT, chemical grafting	hBMSCs	S. aureus and E. coli		The coating exhibits good osseogenesis and antibacterial properties.	[154]
38	Osteogenesis/Antibacterial	BMP-2/CS/HA	MAO, dip coating	MC3T3-E1 cells	E. coli		CS encapsulation sustains BMP-2 release with added antibacterial properties.	[115]
39	Antibacterial	Vancomycin	MAO, dip coating			The rabbit osteomyelitis model (infection with MRSA)	In vivo studies demonstrate the potential of this coating to prevent MRSA infection.	[148]
40	Osteogenesis/Antibacterial	Vancomycin	MAO, dip coating, chemical grafting	BMSCs	S. aureus	Rat femur	Functional coatings prevent prosthesis infection and promote bone integration at the interface.	[155]
41	Antibacterial	Mesoporous silica NPs (MSNs), octenidine (OCT)	Electrophoretic-enhanced MAO	OBs	S. aureus and E. coli		The coating exhibits good osseogenesis and antibacterial properties.	[156]
42	Bioactivity/Antibacterial	N, Bi	MAO, photocatalysis	HGFs	Streptococcus sanguinis and Actinomyces nasseri		The coating has bactericidal properties under visible light.	[123]
43	Osteogenesis/Antibacterial	MoSe2, CS	MAO, electrospinning, photocatalysis	MC3T3-E1 cells	S. mutans	Rat tibia	Adding MoSe2 significantly enhances TiO2 coating photothermal and photodynamic capabilities.	[161]
44	Skin-integration/Antibacterial	β-FeOOH, Fe-TiO2	MAO, HT, photocatalysis	Mouse fibroblasts (L-929)	S. aureus	Mouse skin infection model	The β-FeOOH/FeTiO2 heterojunction prevents bacterial infection under light irradiation.	[108]
45	Osteogenesis/Anti-inflammatory	Ca, Si	MAO	SaOS-2 cells			The coating inhibits inflammation and induces M2 macrophage polarization.	[167]
46	Antibacterial/Immunoregulation	Cu	MAO	RAW 264.7 macrophages, SaOS-2 cells	S. aureus		Cu boosts macrophage-driven osteogenesis and antibacterial activity in biomaterials.	[168]
47	Osteogenesis/Anti-inflammatory	Zn	MAO	RAW264.7 macrophages, BMSCs			The coating shows good osteogenic and anti-inflammatory properties.	[169]
48	Osteogenesis/Anti-inflammatory	Mg	MAO	RAW 264.7 macrophages			Mg acts as an anti-inflammatory agent, inhibiting inflammation and promoting osteogenesis.	[170]
49	Anti-inflammatory	Co	MAO	RAW 264.7 macrophages		Mouse air chamber model	Cobalt-loaded Ti exhibits immune-regulatory effects on macrophages.	[171]
50	Osteogenesis/Angiogenesis/Anti-inflammatory	Li	MAO	BMDMs, mouse embryonic cell line (C3H10T1/2), HUVEC		Mouse air-pouch model	Low Li doses effectively regulate immunity, and promote osteogenesis.	[102]
51	Osteogenesis/Angiogenesis/Anti-inflammatory	HA	MAO, SHT	MC3T3-E1 cells, human umbilical vein fusion cells, RAW 264.7 cells		Rabbit femur	This coating promotes osteogenesis and angiogenesis, and induces M2 macrophage phenotype.	[172]
52	Osteogenesis/Anti-inflammatory	HA	MAO, SHT	MC3T3-E1 cells, endothelial cells, RAW 264.7 cells		Rabbit femur	Nanoparticle-shaped HA is beneficial for osteogenesis, angiogenesis, and immune regulation, whereas nanorod-shaped HA is the opposite.	[173]
53	Osteogenesis/Anti-inflammatory	SiO2, ZnPs	MAO, sol-gel	MC3T3-E1 cells			The coating shows good osteogenic and anti-inflammatory properties.	[175]
54	Osteogenesis/Anti-inflammatory	Sr, silk fibroin-based wogonin NPs	MAO, electrochemical deposition, LBL	RAW 264.7 cells, OBs		Osteoporotic rat femur	The coating shows good osteogenic and anti-inflammatory properties.	[176]

4. Applications and Challenges of MAO in Implants

4.1. Applications

MAO coatings, particularly with specific functional additives, possess remarkable biological functionalities including bone promotion, antimicrobial properties, and anti-inflammatory effects. Though clinical applications of Ti-based implants with MAO-modified functional coatings are limited, animal experiments have validated several implants with such coatings (Table 3). Presently, surface modification techniques like sandblasting, acid etching, and anodization, along with the introduction of bioactive materials such as calcium phosphate, hydroxyapatite, and calcium nanoparticles, are widely employed to optimize implant surface microstructure and topography. These modifications primarily aim to enhance biocompatibility and promote osseointegration. For example, MAO techniques have been utilized to create rough and porous surfaces on dental implants and femoral stems, facilitating improved tissue bonding [119].

TiUnite and TiUltra surfaces(Nobel Biocare, Zurich, Switzerland), along with the Ospol implant(OspolAB, Malmö, Sweden) and the M implant(Shinhung, Seoul, Republic of Korea), are all manufactured using the MAO process and are commercially available. The TiUnite surface series by Nobel Biocare^TM, introduced in the early 2000s, has undergone continuous refinement and gained worldwide acceptance. These implants, containing 7% phosphorus in the form of titanium phosphate chemical bonds, are fabricated using an electrolyte mixture with phosphorus. A comprehensive review [178] encompassing prospective studies spanning from 2000 to 2016 on the clinical performance of implants featuring the TiUnite surface affirms the high survival rates and favorable maintenance of marginal bone associated with such implants. Incidences of peri-implantitis are notably lower for implants employing the TiUnite surface. Notably, implants with this surface consistently exhibit predictable treatment outcomes across diverse indications. In 2019, Nobel Biocare^TM introduced the TiUltra surface series (Figure 9), which features a progressively rough and porous surface texture extending from the implant collar to its tip, exhibiting excellent hydrophilic properties [180]. Moreover, the M implants of Shinhung^TM are produced using the MAO method with an electrolyte mixture containing magnesium, resulting in a TiO₂ surface containing magnesium (≤9.3%) and phosphorus (≤3%) [181]. Similarly, Ospol^TM implants are created using the MAO method with an electrolyte mixture containing Ca, leading to the incorporation of calcium ions (less than 11%) within the TiO₂ in the form of calcium titanium oxide bonds [182].

Femoral shafts treated with sandblasting and MAO coating are commercially available. For example, the Korean Bencox^TM hip system (Corentec, Seoul, Republic of Korea) has cementless sandblasted femoral stems with MAO coating on the surface (Figure 10). Made from Ti-6Al-4V alloy, it is a double-wedge straight tapered stem with a rectangular cross-section. The coating method of MAO coating is as follows: After Ti plating, the sample is electrochemically oxidized using the MAO process. The sample is subjected to MAO treatment with a DC pulse power supply in an aqueous electrolyte containing Ca and P.

In a study by Lim et al. [179], in a follow-up period of at least 5 years, the MAO-coated sandblasted surface did not demonstrate superiority over primary total hip arthroplasty (THA) with the same femoral component design. While MAO is theoretically known to promote bone growth through calcium and phosphorus binding to the titanium alloy surface, forming thick oxides and nano-porous coatings, its clinical utility has not been confirmed. In a study with 10 years of follow up [183], 309 cases of THA (involving 256 patients) were performed using the non-cemented Bencox^TM hip joint system. A Kaplan–Meier survival analysis indicated a 10-year survival rate of 97.4% when using revision for any reason as the endpoint, and a survival rate of 98.7% when using aseptic loosening revision as the endpoint. While there was a decrease in the proportion of cases with bone resorption in the follow up, the absence of a control group precludes concluding the ability of MAO coatings to resist bone resorption, despite various studies suggesting their potential for preventing bone resorption. In summary, after a follow-up period of at least 10 years, the outcomes of non-cemented sandblasted and MAO-coated THA tapered wedge stems are satisfactory.

The MAO coating cannot only promote bone integration but also improve the wear and corrosion resistance of implants. Khanna et al. [184] performed cold spraying to deposit an Al metal layer onto a Ti-6Al-4V alloy substrate, followed by heat treatment to enhance adhesion. They then conducted MAO treatment to form a dense α-Al₂O₃ layer on the substrate surface, with a Vickers hardness matching that of sintered alumina used for femoral heads (Figure 11).

4.2. Challenges

Although MAO coatings have undergone extensive research and led to the emergence of related products, there are still challenges and limitations in the production and application:

• Despite considerable progress in studying the discharge process of MAO, some micro-scale mechanisms remain unclear, such as cathodic discharge and soft spark discharge. This uncertainty affects the control of the microstructure of MAO coatings [22].
• In the coating manufacturing process, establishing the electrolyte composition and electrical parameters still necessitates multiple experiments. Defining the optimal parameters remains challenging. Minor losses of electrolytes during usage and the settling of particulate or colloidal electrolytes can also impact coating performance [44,54].
• The impact of coating morphology on cell bioactivity remains inconclusive. While some studies suggest that moderate roughness aids in cell adhesion and proliferation, and porous surfaces facilitate cell osteogenic differentiation, there is still debate about the optimal pore size. Different cell types may have varied requirements for morphology [90,91]. Additionally, due to cracks and interconnected pores, their durability and wear resistance require attention [20].
• Coatings with added functionalities, such as antimicrobial and anti-inflammatory components, may exhibit toxic effects on normal cells, especially with high concentrations of antibiotics or antimicrobial agents [119,164].
• There is a substantial amount of research on biologically functional coatings; the availability of implants in the market is relatively limited. Successfully translating these technologies into commercial products may face greater challenges [119].

5. Conclusions and Prospects

5.1. Conclusions

The main focus of this paper revolves around a series of factors influencing the performance of MAO coatings and the biological functions of these coatings under various treatment conditions. As a result, the following conclusions are drawn:

• The discharge process is pivotal, influenced by the electrolyte, electrical parameters, and substrate. The electrolyte affects the composition and structure of coatings. Electrical parameters impact coating performance by altering discharge energy. Additionally, the substrate and pre-treatment also contribute.
• MAO transforms implant surfaces, creating varied topologies and infusing substances for osteointegration. It provides multiple strategies to combat surface infections by adding antibacterial agents. Recent surface engineering maximizes inflammation benefits while minimizing risks, with MAO coatings showing considerable potential. The biocompatibility and long-term stability of coatings require more attention.
• Presently, MAO is being clinically applied to promote bone integration, especially in the field of dentistry. Despite some progress in the MAO bioactive coatings on Ti and its alloys, challenges still exist in the production and application of coatings.

5.2. Prospects

The progress and challenges of MAO coatings suggest considerable research and application potential for implant surfaces. The following outlines prospects:

• Certain micro-mechanisms in the MAO process need more study. Collaborative efforts across disciplines will enhance the understanding of discharge, electrolyte behavior, and coating growth complexities. Real-time detection, high-resolution characterization, and numerical simulation may unveil the link between microstructures and performance, paving the way for tailored and controllable coatings.
• Developments in electrolytes will prioritize optimizing formulations, extending lifespan, and emphasizing sustainability by using renewable or recyclable components. Equipment improvements will focus on enhancing power supply design for increased efficiency and stability. Orthogonal experiments can be used to validate new formulations and equipment designs under various conditions. Standardized and automated manufacturing will enhance production efficiency and reduce waste, contributing to the sustainability and cost-effectiveness of MAO technology.
• The integration of biology, materials science, and engineering will be employed to design a controllable coating. Comprehensive experiments are imperative to unveil the influence of morphology on cellular behavior. Incorporating nanoparticles into the electrolyte or using other techniques to repair defects in MAO coatings could potentially enhance wear resistance and durability.
• Optimizing antimicrobial and anti-inflammatory components in coatings is crucial. Researchers need to delve into understanding mechanisms of action to reduce concentrations while maintaining efficiency and minimizing adverse effects. Avoiding overuse aims to slow bacterial resistance. When designing coatings, selecting friendly components and comprehensive biocompatibility testing will ensure coatings do not trigger allergic or inflammatory responses.
• MAO coatings have vast market opportunities. However, successful market entry requires meeting strict approval and regulatory standards. Key factors include in-depth research on biocompatibility, drug release kinetics, and rigorous safety testing. The evaluation methods and standards for coatings also need to be further developed. Collaboration, innovation, and a continued focus on product safety will drive the successful commercialization of coatings in the future.