Microstructure and Properties of Biomedical Mg-Zn-Ca-Ag Alloy and the Micro-Arc Oxidation Coatings

Abstract

This study investigates the influence of Ag addition on the microstructure, mechanical behavior, corrosion resistance, and antibacterial performance of Mg-Zn-Ca-Ag alloys and their micro-arc oxidation (MAO) coatings. Four casting alloys containing 0.2, 0.4, 0,6 and 0.8 wt.% Ag were fabricated and characterized by SEM, XRD, and TEM. The microstructure consisted mainly of α-Mg, Mg₂Ca, Mg₇Zn₃, and Mg₆Ca₂Zn₃ phases, and the elastic modulus (~25.8 GPa) was comparable to that of human bone. MAO coatings produced in a bio-functional electrolyte exhibited pit-like morphologies due to Ag-induced melt fluidity and self-sealing effects. The coatings were composed primarily of MgO, Mg₂SiO₄, Ca₃(PO₄)₂, CaCO₃, and Ag₂O, with the ZQ 0.8-MAO sample showing the highest Ca/P ratio (1.75), indicative of superior bioactivity. Electrochemical impedance spectroscopy revealed optimal corrosion resistance (2.56 × 10⁴ Ω·cm²), while antibacterial efficiency exceeded 96%. Overall, Ag alloying enhanced both the bulk and surface properties of Mg-Zn-Ca alloys, yielding robust, corrosion-resistant, and antibacterial coatings with excellent biocompatibility-highlighting their potential for biodegradable orthopedic implant applications.

1. Introduction

Magnesium (Mg) is one of the most abundant elements in the Earth and an essential trace element in the human body [1]. It participates in most intracellular biochemical reactions and ranks second only to potassium ions in cellular concentration. When magnesium-based implants are introduced into the body, they can degrade in situ, releasing Mg²⁺ ions that contribute to protein synthesis, act as cofactors for enzymatic reactions, assist in signal transduction, and promote proper muscle and bone contraction [2]. These characteristics make magnesium a desirable candidate for use in temporary biomedical implants. However, magnesium’s low standard electrode potential (−2.73 V) leads to rapid electrochemical degradation in aqueous environments, such as body fluids [3,4]. As a result, magnesium alloys often fail to maintain structural integrity long enough to support tissue healing, which significantly limits their clinical application as biodegradable implants.

To address these limitations, researchers have widely explored alloying as a method to enhance the corrosion resistance, mechanical strength, and biocompatibility of magnesium materials. Common alloying elements include aluminum (Al), zinc (Zn), calcium (Ca), strontium (Sr), and silver (Ag), resulting in binary and ternary systems such as Mg-Ca, Mg-Zn, Mg-Sr, and Mg-Zn-X (where X = Zr, Mn, Sr, Ca, etc.) [5,6,7,8,9,10,11,12] Among these, Zn and Ca are frequently used due to their biocompatibility and structural roles in bone metabolism. Silver improves magnesium alloys by refining the grain structure, increasing tensile strength and ductility, and enhancing corrosion resistance in some conditions. Zn contributes to corrosion resistance and mechanical strength, while Ca refines grain size and promotes the formation of biologically favorable secondary phases. Ag has attracted considerable attention as an alloying element owing to its pronounced and wide-spectrum antibacterial activity, which remains effective even at trace levels. It suppresses bacterial proliferation primarily by compromising cell membranes, interfering with essential enzyme systems, and inducing DNA damage [13]. In addition to its potent antibacterial action, incorporating Ag into magnesium alloys enhances their corrosion resistance and mechanical integrity. Previous studies have demonstrated that when the Ag content is maintained below 1 wt.%, it exerts minimal cytotoxicity and does not significantly hinder cellular adhesion or proliferation [14]. However, it was also reported that Mg-Ag phases can significantly accelerate the corrosion process of Mg-Ag alloys whereas Ag in solid solution may slightly affect the corrosion rate of Mg-Ag alloys under simulated physiological condition [15].

Although alloying can effectively improve the properties of magnesium-based alloys, their corrosion resistance is still insufficient to meet the clinical service requirements of biodegradable implants, which typically range between 90 and 180 days. Consequently, further surface modification strategies are essential to prolong the functional lifespan of magnesium implants and ensure their reliable in vivo performance. Among the various surface modification techniques, micro-arc oxidation (MAO)-also known as plasma electrolytic oxidation-has emerged as one of the most promising approaches. Through electrochemical plasma discharges in an alkaline electrolyte, this process generates a dense, firmly adherent, and ceramic-like oxide coating on magnesium and other valve metals, providing excellent bonding strength and structural stability [16,17,18]. Compared to other coating methods, MAO offers several advantages [19,20,21], MAO coating possesses Strong adhesive force, excellent corrosion resistance, porous tunable composition, simple and controllable processing. These unique characteristics have enabled micro-arc oxidation coatings to find widespread use in the automotive, aerospace, and biomedical industries. When applied to magnesium alloys, MAO effectively improves corrosion resistance while simultaneously enhancing bioactivity, primarily through the incorporation of Ca, P, and Si species from the electrolyte into the oxide coating. Nevertheless, the intentional addition of functional elements such as Ag into MAO coatings has received little attention, particularly in quaternary Mg-Zn-Ca-Ag alloy systems.

Previous studies have primarily focused on binary or ternary Mg-based systems. For example, Mahmood Razzaghi et al. [13] investigated the effects of Zn and Ag on corrosion behavior and antibacterial performance in Mg-3Zn-0.5Ag alloys. Although positive results were reported, a comprehensive understanding of how Ag content affects the structural and functional evolution of both the substrate alloy and the MAO coating is still lacking. To bridge this knowledge gap, a novel series of Mg-Zn-Ca-Ag casting alloys containing 0.2, 0.4, 0.6, and 0.8 wt.% Ag-designated as ZQ 0.2, ZQ 0.4, ZQ 0.6, and ZQ 0.8-were developed in this study. The as-cast alloys were comprehensively characterized in terms of their phase constitution, microstructure, mechanical performance, corrosion behavior, surface wettability, and antibacterial activity. Furthermore, the various alloys where coated using the MAO process by using a biologically optimized electrolyte. The resulting coatings were evaluated with respect to their phase composition, surface morphology, elemental distribution, corrosion and wear resistance, interfacial adhesion, wettability, and antibacterial effectiveness.

The present study aims to elucidate the dual function of silver (Ag) as both an alloying constituent and a biofunctional agent. Particular emphasis is placed on determining whether Ag incorporated into the substrate can be effectively transferred into the MAO coating, how its presence influences the evolution of the characteristic “pit-like” surface morphology, and to what extent it contributes to enhanced corrosion resistance and antibacterial performance. The insights obtained are expected to offer a theoretical basis for designing multifunctional, biodegradable magnesium-based implants with bone-compatible mechanical properties, pronounced antibacterial activity, and durable corrosion protection.

2. Materials and Methods

2.1. Alloy Casting

In this study, ZQ casting alloys were prepared using high-purity raw materials. Beijing Yanbang New Material (Beijing, China) supplied pure magnesium (99.99%), while zinc (99.99%) was obtained from (Haima Numerical Control Equipment Co., Ltd., Suzhou, China) Silver granules and the Mg-Ca master alloy, both with 99.99% purity, were also used.

First, a metallographic cutting machine was used to divide the magnesium ingot into smaller blocks weighing approximately 1 kg each. A grinding wheel was then employed to polish their surfaces, removing oxide coatings and exposing the metallic luster. Zinc and Mg-Ca alloy pieces were cut into small segments using wire electrical discharge machining (EDM), followed by surface polishing with sandpaper to eliminate oxidation. After preparation, the raw materials were weighed precisely based on the target alloy compositions and set aside for melting.

During the smelting process, both the mold and metallic ingots were preheated to 200 °C in a resistance furnace for 1 h. Following preheating, the crucible was cleaned, and a mix of sulfur hexafluoride (SF₆) and Carbon dioxide (CO₂) was introduced to provide an inert protective atmosphere. The furnace was then heated to 700 °C to melt the magnesium ingots, followed by an increase to 770 °C to ensure complete liquefaction. The melt was held at this temperature for 20 min. Subsequently, preheated zinc, silver, and Mg-Ca alloy were added to the molten magnesium. The mixture was stirred with a stainless-steel rod while maintaining a fixed contact point, which minimized disturbance of the surface oxide coating and prevented compositional inconsistencies. After homogenization, the melt was held for an additional 20 min at 770 °C.

Before casting, the gas mix was again introduced into the preheated mold to displace residual air. The molten alloy was then poured into the mold, with a stainless-steel rod used to shield the melt from surface oxides. Once cooled to approximately 200 °C, the mold was opened and the ingots were extracted. This casting procedure yielded four ZQ alloys with varying silver contents: ZQ 0.2, ZQ 0.4, ZQ 0.6, and ZQ 0.8. Their nominal compositions are listed in

Table 1. The compositions of ZQ alloys.

Samples	Zn (wt.%)	Ca (wt.%)	Ag (wt.%)	Mg (wt.%)
ZQ 0.2	4.0	0.8	0.2	Balance
ZQ 0.4	4.0	0.6	0.4	Balance
ZQ 0.6	4.0	0.4	0.6	Balance
ZQ 0.8	4.0	0.2	0.8	Balance

.

2.2. MAO Coating Preparation

A pre-optimized biological electrolyte was employed to fabricate the MAO coatings. The electrolyte composition included 0.8 g/L sodium hexametaphosphate ((NaPO₃)₆), 6 g/L sodium metasilicate nonahydrate (Na₂SiO₃·9H₂O), 0.5 g/L calcium acetate monohydrate (Ca(CH₃COO)₂·H₂O), and 0.5 g/L sodium dihydrogen phosphate monohydrate (NaH₂PO₄·H₂O). Sodium hydroxide (NaOH) was added to adjust the pH of the solution to approximately 13.

MAO processing was carried out using a WHD-30 bipolar AC pulse power supply. A stainless-steel tank served as the cathode, while the magnesium alloy sample acted as the anode. The electrolyte was maintained at a constant temperature of ~30 °C by means of an external water circulation system combined with an industrial chiller (ZX-03AC Guangdong Zhongkai Intelligent Equipment Co., Ltd., Dongguan, China).

The coating process utilized a hybrid power control mode involving both constant-voltage and constant-current phases. Electrical parameters were based on prior optimization by the research team. In the constant-voltage stage, the forward and reverse voltages were set at 400 V and 30 V, respectively, with a forward duty cycle of 30% and a reverse duty cycle of 70%, operating at a frequency of 600 Hz. The voltage-controlled stage lasted from 0 to 5 min. Following this, the process transitioned to constant-current control, with forward and reverse currents of 1.2 A and 0.3 A, respectively. The duty cycles remained the same, while the system operated at 400 V and a frequency of 600 Hz. The duration of this stage ranged from 5 to 15 min.

2.3. Characterizations

2.3.1. Microstructure

A Zeiss Merlin (JSM-6480, JEOL, Tokyo, Japan) Compact field emission scanning electron microscope (SEM) was used to characterize the microstructure of the prepared magnesium alloy and the MAO coating. The SEM is equipped with an energy-dispersive spectrometer (EDS) for conducting area and point scan analyses to determine the elemental composition of the alloy and the coating.

A JEM-2100F (JEOL, Tokyo, Japan) field emission transmission electron microscope (TEM) was employed to analyze the microstructure of ZQ alloys. The samples are firstly grounded with SiC paper, then use the Gatan 61 ion thinning instrument to obtain a sample with a thickness of approximately 100 nm. The TEM was operated at an acceleration voltage of 200 kV. The lattice spacing in high-resolution transmission electron microscopy (HRTEM) images was measured using Digital Micrograph (DM 3.6.1) software. Fourier Transform (FFT) of the high-resolution images was performed to measure the distance between corresponding diffraction spots L (nm) L (nm) L (nm). To minimize errors, five measurements were taken, and the average value was used to calculate the lattice plane spacing d = L/2.

For phase analysis of the magnesium alloy and the MAO coating, a Shimadzu XRD-6000 X-ray diffractometer (Bruker AXS GmbH, Karlsruhe, Germany) was used. The phase identification was performed by comparing the three strongest peaks with PDF cards using Jade 9 software. The specific parameters were set as follows: a scanning speed of 2°/min, using Cu Kα radiation with an accelerating voltage of 40 kV, a scanning range of 20° to 80°, and one scan repetition.

2.3.2. Mechanical Properties

Nano-scratch tests were performed using a Swiss-made CSM-NHT2 nano-indentation tester (CSM Tech Canada, Montreal, QC, Canada). The tests utilized a Rockwell indenter with a 100 μm tip radius. The mechanical parameters, such as load, acoustic emission signals, and friction coefficient, were recorded as the indenter penetrated through the MAO coating to the magnesium alloy substrate. This characterized the adhesion strength between the MAO coating and the magnesium alloy substrate. The tests were conducted with a progressive load, with the maximum load set at 15 N and the loading force applied at a rate of 5 N/mm across a 3 mm linear distance on the MAO coating surface.

A CMT5205 universal testing (MTS Systems Co., Ltd., Shenzhen, China) machine was used to perform mechanical tensile strength tests on the cast magnesium alloy. The tensile specimens are round ASTEM E8 samples, and the tensile tests were conducted at a displacement speed of 2 mm/min, a maximum load of 15 N, and a testing length of 3 mm.

A UMT-2 high-temperature tribometer (Center for Tribology Co., Ltd., Campbell, CA, USA) was used to test the wear resistance of the MAO coating. The tests employed a linear friction method, comparing the friction coefficients of different MAO coatings under the same load conditions. The counter face used in the experiment was a steel ball with a surface hardness of HRC63. The testing parameters included a load of 5 N, a friction speed of 50 mm/s, a wear track length of 10 mm, and the test duration is 600 s. After the test, the wear tracks were wiped clean with alcohol, and the wear morphology and wear area were observed using a LEXT OLS4000 confocal laser microscope from Olympus, Tokyo, Japan.

2.3.3. Electrochemical Analyses

A CHI660E electrochemical workstation, manufactured by Shanghai Chenhua Instrument Co., Ltd. (Shanghai, China), was used to measure the impedance of the magnesium alloy and the MAO coating. The electrochemical testing was conducted using a three-electrode system, with a saturated calomel electrode (SCE) as the reference electrode, a platinum electrode as the auxiliary electrode, and the sample under test as the working electrode. The testing environment was Hank’s solution maintained at 37 °C. After testing, the corrosion resistance of the alloy and the coating in Hank’s solution was analyzed using ZsimpWin v3.60 and Zview2 v3.2.2.22 software for equivalent circuit fitting. The potentiodynamic polarization test was carried out after EIS test with a scanning range of −200 mV~2 V (vs. OCP) and a scanning rate of 0.001 mV/s. The software Cview2.0 was used for the potentiodynamic polarization test.

2.3.4. Antibacterial Tests

The antibacterial tests were realized using the following steps: First, the preparation steps. The prepared liquid culture media are bottled and sterilized using autoclaving at 121 °C for 30 min. The solid culture media are also placed in an autoclave and sterilized under the same conditions. After sterilization, the solid media are poured into Petri dishes while still warm and allowed to cool and solidify. The activation method for Escherichia coli (E. coli) is as follows: A specific amount of lyophilized bacteria is added to the prepared liquid culture medium, which has been cooled to room temperature. This mixture is then placed in an incubator shaker set at 37 °C and 100 rpm, where it is cultured for 24 h.

Then, the experimental steps. Turn on the UV-Vis spectrophotometer and allow it to warm up for 20 min. Place a measured amount of liquid culture medium into a glass cuvette to serve as the control. Take a small amount of the E. coli culture, place it into an Erlenmeyer flask, and slowly dilute it with distilled water to a specific concentration. Then, a measured amount of the diluted bacterial suspension is transferred into another glass cuvette. Direct the light beam at the control cuvette, press the “Zero” button to set the absorbance to zero, and then direct the light beam at the cuvette containing the bacterial suspension. Observe the absorbance value. When the wavelength is set to 600 nm, and the absorbance value is approximately 0.014, the bacterial suspension concentration can be estimated at around 10⁵ cfu/mL. If the absorbance differs significantly, adjust the amounts of bacterial suspension or distilled water and repeat the dilution and absorbance measurement.

Take a measured amount of the bacterial suspension, with a concentration of approximately 10⁵ cfu/mL, and add it to a new liquid culture medium in an Erlenmeyer flask. Place the pre-sterilized magnesium alloy samples of different compositions into the flasks, label each flask, cover with sealing film, and incubate them in a shaking incubator at 37 °C and 100 rpm for 6 h. After incubation, two 1 mL samples of the culture medium were taken and placed into two 1.5 mL centrifuge tubes. Centrifuge the samples at 4000 r/min for 1 min. Remove 0.5 mL of the supernatant from each centrifuge tube, combine the remaining culture media of the same sample type, and transfer the mixture into a 15 mL centrifuge tube containing 9 mL of sterile 0.9% sodium chloride solution. Vortex the tube for 1 min. Repeat this step to obtain bacterial suspensions with different dilution factors, adjusting the dilution factor according to the specific experimental requirements. Spread 200 μL of the culture medium onto solid culture media plates and incubate them in a 37 °C incubator for 10 h. Observe the growth of bacterial colonies on the plates, the biological test is performed with 3 parallel samples for each kind.

2.3.5. Wettability Test

A JC2001D1 contact angle goniometer (Shanghai auto-lab technology Co., Ltd., Shanghai, China).was employed to measure the water contact angle of the magnesium alloy and the MAO coating. The static drop method was used, where the contact angle between the water droplet edge and the surface of the coating was measured to assess the hydrophilicity of the alloy and the coating. The contact angle was calculated by software, which processed the dynamic images of the water droplet captured by a camera.

3. Results and Discussion

3.1. Properties of ZQ Alloys

In this section, ZQ alloys are studied in elemental composition, phase composition, mechanical properties, corrosion resistance, wettability, and antibacterial capacity. ZQ alloys are comprehensively evaluated as a biomedical material.

3.1.1. Phase Composition and Elemental Composition of ZQ Alloys

Figure 1 presents the X-ray diffraction (XRD) patterns of ZQ alloys with varying compositions. Based on the Mg-Ca binary phase diagram and corresponding XRD spectra, the primary phases identified in the alloys include α-Mg, Mg₂Ca, Mg₆Ca₂Zn₃, Mg₇Zn₃ (also denoted as Mg₅₁Zn₂₀), and γ-Ag. The data indicate that the presence of the Mg₂Ca phase diminishes as the Ca content in the alloy decreases. This trend is consistent with findings reported by Larionova et al., who demonstrated that the formation of secondary phases in Mg-Zn-Ca systems is governed by the Zn/Ca atomic ratio [22]. Specifically, when the Zn/Ca ratio exceeds 1.4, the resulting phases typically include only α-Mg and Mg₆Ca₂Zn₃ [23,24].

In the present study, the alloy with the lowest Zn/Ca ratio is ZQ 0.2, with a value of 3.06. According to theoretical predictions based on the binary phase diagram, this composition should yield only α-Mg and Mg₆Ca₂Zn₃ as the secondary phases. However, XRD analysis reveals the unexpected presence of Mg₂Ca. This can be explained by the solidification behavior of the alloy, α-Mg nucleates first during cooling, and due to its smaller atomic radius and lower diffusion activation energy, Zn diffuses more rapidly into the remaining liquid than Ca. This preferential Zn enrichment suppresses Ca accumulation at the solid–liquid interface. As solidification proceeds and α-Mg grains grow, the trapped Ca atoms within the grains reach supersaturation, ultimately leading to the precipitation of Mg₂Ca upon further cooling [25]. When the overall Ca content is further reduced, this supersaturation condition is no longer met, thereby suppressing the formation of Mg₂Ca during solidification [26]. Mg₂Ca has been reported to promote grain refinement in magnesium alloys, which is generally beneficial for mechanical performance. However, an excessive Ca content can lead to the formation of abundant secondary phases that exacerbate micro-galvanic corrosion, thereby compromising corrosion resistance [27]. For this reason, it is critical to regulate the Ca concentration within the alloy. Previous studies have shown that maintaining Ca levels below 1 wt.% yields a favorable balance between mechanical properties and corrosion resistance [28].

Furthermore, analysis of the Mg-Zn binary phase diagram reveals that a eutectic reaction occurs at 340 °C, producing α-Mg and Mg₇Zn₃ phases. Mg₇Zn₃ is metastable and typically undergoes eutectoid transformation to form the thermodynamically stable MgZn phase at approximately 325 °C. The absence of MgZn peaks in the XRD patterns suggests that the cooling rate during alloy solidification was sufficiently high to suppress this transformation, resulting in the retention of the metastable Mg₇Zn₃ phase [29].

Figure 2 represents the TEM results of ZQ alloys. Figure 2a is the diffraction pattern of Ca₂Mg₆Zn₃ of crystal indices of (2 2 0), the interplanar crystal spacing d = 2.44 Å. Figure 2b represents the diffraction pattern of Mg₂Ca (2 0 1), d = 2.53 Å. Figure 2c shows the diffraction pattern of Mg₇Zn₃, the crystal indices are not shown in PDF card, yet d = 2.30 Å, The TEM results are coherent with the XRD spectrums.

3.1.2. Microscopic Morphology of ZQ Alloys

Figure 3 illustrates the microstructural evolution of the ZQ alloys, highlighting the distribution and morphology of secondary phases. According to previous studies, the dotted precipitates correspond to the Mg₂Ca phase, while the strip-like features are primarily identified as Mg₆Ca₂Zn₃ [30].

In Figure 3a, the ZQ 0.2 alloy exhibits a microstructure characterized predominantly by randomly distributed, dot-shaped precipitates, with fewer strip-like features. These precipitates are clearly distinguishable on the alloy surface, having formed during the solidification process. As the matrix composition changes across the alloy series, the morphology of the precipitates transitions from a dispersed dot-like distribution to a more organized arrangement, marked by a relative decrease in Mg₂Ca and an increase in elongated Mg₆Ca₂Zn₃ structures. In Figure 3b, the microstructure of the ZQ 0.4 alloy reveals that the previously random dot-like precipitates have begun to align preferentially along the grain boundaries, effectively delineating the grain morphology. The accumulation of precipitates at grain boundaries serves to hinder grain coarsening, thereby contributing to the enhancement of the alloy’s mechanical properties. Figure 3c presents the microstructure of the ZQ 0.6 alloy, although the quantity of strip-like precipitates is comparable to that in the ZQ 0.4 alloy, the overall distribution is more random, and the grain boundaries are less discernible. The Mg₆Ca₂Zn₃ phase exhibits a relatively small potential difference with the α-Mg matrix, whereas Mg₂Ca, with its larger potential difference, is more susceptible to micro-galvanic corrosion. Consequently, an increased proportion of strip-like precipitates may contribute to a modest improvement in corrosion resistance. In Figure 3d, the ZQ 0.8 alloy displays a further reduction in Mg₂Ca precipitates, with the majority of secondary phases appearing as uniformly distributed, strip-like features, the element mapping of ZQ 0.8 alloy in Figure 3e also confirms it. This morphology effectively inhibits grain boundary migration and promotes grain refinement, thereby enhancing the alloy’s mechanical strength. Additionally, the predominance of Mg₆Ca₂Zn₃ contributes to an incremental improvement in corrosion resistance due to the reduced electrochemical potential mismatch with the surrounding matrix.

3.1.3. Mechanical Properties of ZQ Alloys

Figure 4 presents the hardness and tensile strength data for the ZQ alloys with varying compositions. The changes in Ca and Ag content within the matrix significantly influence the type, quantity, and spatial distribution of precipitates. These microstructural variations in turn lead to measurable differences in the mechanical performance of the alloys.

Conventional biomedical metals such as Ti-6Al-4V, WE43, and 316L stainless steel exhibit considerable mismatches in both elastic modulus and tensile strength when compared to human bone, for example, the tensile strength of Ti-6Al-4V is 860–965 MPa, the tensile strength of WE43 is 220 MPa, while that of human bone tissue is 1.5–38 MPa [3,31,32]. These disparities often result in stress shielding, which can impair bone regeneration. In contrast, the magnesium alloys developed in this study demonstrate elastic moduli much closer to that of natural bone. The ZQ 0.6 and ZQ 0.8 alloys exhibit elastic moduli in the range of 25–30 GPa, aligning well with the mechanical behavior of human cortical bone. Their tensile strengths also fall within acceptable limits for load-bearing orthopedic applications. These results indicate that the newly developed alloys offer mechanical properties more compatible with biological tissues than those of traditional implant metals. The enhanced performance can be attributed to the formation of finely distributed intermetallic compounds during solidification, resulting from the controlled addition of alloying elements. These intermetallic effectively impede grain boundary migration, thereby contributing to improved mechanical strength. However, variations in the distribution and morphology of these phases among different alloy compositions lead to slight differences in their overall mechanical performance.

3.1.4. Potentiodynamic Polarization

Figure 5 displays the electrochemical testing results for the four ZQ alloy compositions, while

Table 2. Fitting data of alloys with different compositions.

Alloys	Corrosion Current Density (A·cm2)	Corrosion Potential (V)
ZQ 0.2	4.70 × 10−4	−1.42
ZQ 0.4	5.46 × 10−4	−1.45
ZQ 0.6	5.39 × 10−4	−1.41
ZQ 0.8	1.87 × 10−4	−1.40

summarizes the corresponding polarization curve fitting data. According to electrochemical theory, corrosion in magnesium alloys tends to initiate in the vicinity of secondary phases, irrespective of whether these phases are distributed as isolated particles or form continuous networks along grain boundaries. For alloys with relatively low Zn and Ca content, the dominant secondary phases include Mg₂Ca, α-Mg, and Mg₆Ca₂Zn₃. The standard electrode potentials of these phases follow the order: Mg₂Ca < α-Mg < Mg₆Ca₂Zn₃ [33,34]. The resulting potential differences promote micro-galvanic corrosion, with Mg₂Ca acting as the anodic site and thus being preferentially corroded.

Studies have shown that the addition of Mn, Nd, and other elements to Mg alloys does not cause significant cathodic activation behavior, whereas the addition of Ag to Mg alloys leads to notable cathodic activation behavior. An excessive amount of Ag ions on the metal surface results in a decrease in the corrosion resistance of the alloy, and corrosion mainly occurs in Ag-rich regions and Mg-Ag intermetallic phase regions. The standard electrode potentials of Mg and Ag are −2.37 V and 0.8 V (SHE), respectively. When a galvanic couple is formed between Mg and Ag in the alloy, Mg dissolves preferentially, whether it is in the Mg matrix or the Mg-Ag intermetallic phase [35]. Once silver atoms are separated from the alloy, they are oxidized to Ag⁺ ions, and these Ag⁺ ions react with Mg through a displacement reaction, causing rapid localized corrosion. When Ag is added to the Mg-Zn alloy, the corrosion rate of the alloy increases with the increase in the amount of Ag added. However, when the amount of Ag added is relatively small, the corrosion rate of the alloy is roughly the same as that of the alloy without Ag added. It can be known from the phase analysis in Figure 1 that no Mg-Ag intermetallic compounds are formed in the alloy discussed in this study, and the content of Ag added in this study is relatively low; therefore, its impact on the corrosion resistance of the alloy is minor [36].

As shown in Figure 3, the ZQ 0.8 alloy contains fewer Mg₂Ca precipitates, which reduces the risk of galvanic interactions and enhances its overall corrosion resistance. In contrast, although the ZQ 0.4 alloy still contains Mg₂Ca, the presence of the Mg₆Ca₂Zn₃ ternary phase along grain boundaries appears to play a moderating role during the corrosion process. This phase may temporarily inhibit localized attack by redirecting the corrosion pathway away from the α-Mg matrix and Mg₂Ca, thereby reducing the overall degradation rate [37]. Consequently, both ZQ 0.4 and ZQ 0.8 alloys exhibit comparable corrosion resistance in electrochemical testing.

The polarization curve analysis further indicates that alloying modifications do not significantly alter the self-corrosion current density among the ZQ alloy series. The most favorable self-corrosion potential recorded is −1.40 V, which remains substantially more noble than the standard electrode potential of pure magnesium (−2.37 V). Despite these improvements, magnesium alloys remain inherently prone to corrosion. Therefore, further surface modification strategies, such as coating or anodization, are necessary to achieve enhanced long-term corrosion protection.

3.1.5. Antibacterial Capacity and Wettability of ZQ Alloys

Figure 6 presents the antibacterial performance of the four ZQ alloys. The results clearly demonstrate that the incorporation of silver significantly enhances the antibacterial efficacy of the materials. Compared to the control group, all four alloys exhibit a marked reduction in bacterial colony formation, with antibacterial rates exceeding 97%. These values substantially surpass the 90% threshold defined by the National Health Bureau for antibacterial materials. Notably, the ZQ 0.8 alloy achieves an antibacterial rate of 99.1%, highlighting the excellent antimicrobial performance of the optimized composition. As reported by Lin et al. [38], the implantation of magnesium alloys is typically accompanied by a sharp increase in local pH due to rapid corrosion. Escherichia coli, which proliferates optimally within a pH range of 6–8 [39,40], is highly sensitive to such alkaline conditions, contributing to the initial antibacterial effect of the material shortly after implantation. Furthermore, the metabolic by-products released by E. coli accelerate the corrosion of the magnesium alloy, creating a feedback loop that enhances ion release. As the alloy degrades, Ag⁺ are continuously released into the surrounding environment. These ions bind to bacterial cell wall proteins, penetrate the cytoplasm, and interfere with nucleic acid structures, ultimately disrupting DNA and RNA function and inducing programmed apoptosis [41]. This dual mechanism-alkaline microenvironment and sustained Ag⁺ release, ensures both immediate and long-term antibacterial activity, making the newly developed magnesium-silver alloys highly promising for biomedical implant applications.

Wettability is a critical parameter for evaluating the biocompatibility of biomaterials. It has been well established that surfaces with higher hydrophilicity promote cell adhesion, spreading, and proliferation, thereby enhancing their integration with biological tissues. Typically, materials with a water contact angle less than 80° are considered to possess good biocompatibility due to their favorable surface energy characteristics [37,42,43]. Figure 6f presents the results of wettability and antibacterial tests for the four ZQ alloy compositions. Among these, the highest measured water contact angle is 103.35°, and the lowest is 98.69°, indicating that all samples exhibit relatively high hydrophobicity, with contact angles clustered around 100°. Such hydrophobic surfaces hinder the stable adhesion of cells, particularly under dynamic physiological conditions where fluid movement can displace loosely attached cells. The poor wettability implies limited initial cell anchorage, which is a critical requirement for successful implant integration.

Therefore, despite the favorable mechanical and corrosion properties of ZQ alloys, its high contact angles suggest poor inherent biocompatibility. Without MAO bioactive coating, or plasma treatment, these magnesium alloys are unsuitable for direct application as biomedical implants.

3.2. Properties of ZQ 0.8-MAO Coating

Modern biomedical implants demand degradable materials that satisfy several essential criteria, the ions released during degradation must be non-toxic to surrounding tissues, the material must provide adequate corrosion resistance throughout the healing period, its mechanical strength and elastic modulus should closely match those of natural bone, and it should exhibit reliable antibacterial activity during the entire implantation process. The ZQ alloy developed in this study demonstrates promising characteristics in line with these requirements. Compared to conventional commercial magnesium alloys, the new composition exhibits mechanical properties more closely aligned with those of human bone, as well as favorable biocompatibility and strong antibacterial performance. However, despite these advantages, alloying alone is insufficient to confer the level of corrosion resistance required for clinical applications. To address this limitation, surface modification techniques such as MAO are essential. MAO treatment forms a protective ceramic-like coating on the alloy surface, significantly enhancing its corrosion resistance without compromising its biodegradability or antibacterial function. Therefore, the integration of alloy design with optimized surface treatment is critical for developing next-generation biodegradable implants suitable for orthopedic and biomedical use.

In this section, ZQ 0.8-MAO coating is prepared on the ZQ alloys in paralleling groups with same experimental conditions. The ZQ 0.8-MAO samples are tested and studied in microstructure, phase and elemental composition, mechanical properties, corrosion resistance, wettability, and antibacterial capacity.

3.2.1. Microstructure of ZQ 0.8-MAO Coating

Figure 7 presents the surface morphology and Cross-Sectional Morphology of the ZQ 0.8-MAO coating. The images reveal the presence of characteristic features such as pores, cracks, and accumulated molten material. These molten accumulations result from high-energy micro-arc discharges that occur during the MAO process, where localized high temperatures and pressures cause substrate material to melt, react with the electrolyte, and be ejected through discharge channels. Upon contact with the cooler electrolyte, the molten material rapidly solidifies and adheres to the coating surface. During the constant-current phase of the MAO process, while the current remains stable, the voltage gradually increases. This triggers a repetitive cycle of coating formation, dielectric breakdown, ejection of molten material, re-solidification and reformation of the coating, ultimately leading to the development of a relatively compact ceramic-like oxide coating.

Figure 7b shows the cross-sectional morphology of ZQ 0.8-MAO sample, pores and cracks are observed. These localized defects reduce coating thickness and serve as vulnerable points for corrosion initiation, highlighting the need for further optimization of the MAO process parameters to enhance coating uniformity and barrier performance.

Elemental analysis was conducted at different positions on the surface of the ZQ 0.8 MAO coating. The results show that the elemental content varies significantly among micropores, the edges of micropores, and regions without micropores. As shown in Figure 7c, there are no micropores on the coating surface at this location, elemental analysis reveals that the contents of Ca and Ag elements are extremely low here. However, the contents of Ca and Ag elements begin to increase at the edges of micropores, the Ca content increases from 0.7 to 15.4, and the Ag content increases from 0.16 to 1.66 (as shown in Figure 7d). At the center of the micropores, the Ca content increases sharply to 38.68, while the Ag content rises to 2.26 (as shown in Figure 7e). The reason why the content of Ca is higher than that of Ag is mainly that Ca can undergo chemical reactions not only from the matrix but also from the electrolyte, whereas Ag exists only in the substrate, thus its content is lower than that of Ca.

During micro-arc oxidation, the high applied voltage causes dielectric breakdown of the passive oxide coating formed in the initial anodic phase, resulting in the formation of localized discharge channels. At these sites, intense localized heating leads to melting of the substrate and subsequent chemical reactions with the electrolyte. The molten material is ejected through the discharge channels and rapidly solidifies upon reaching the cooler coating surface, a typical “crater-like” micro-arc oxidation morphology as shown in Figure 7f is formed. Precisely due to this discharge characteristic of micro-arc oxidation, Ca²⁺ and Ag⁺ ions in the substrate are ejected onto the surface of the coating through the discharge channels during the micro-arc oxidation process. Therefore, the contents of these two elements are relatively high in and around the micropores (as shown in Figure 7h,i). In contrast, in areas without micropores, the surface is covered by molten substances (mainly composed of oxides such as MgO), which prevents Ca²⁺ and Ag⁺ from depositing on the surface. As a result, their contents in these areas are extremely low, with Ag⁺ in particular being barely detectable.

The Ca/P ratio on the surface of the MAO coating prepared from the ZQ 0.8 alloy is 1.75, which is close to the Ca/P ratio of human bone (1.67). Therefore, this coating can be considered to have good biological properties and is capable of inducing the growth of osteocytes on its surface.

3.2.2. Phases and Elemental Compositions of MAO Coating on the ZQ Alloys

Figure 8a presents the XRD spectra of the MAO coatings formed on the ZQ alloys, while Figure 8b–d displays the corresponding XPS results. The XRD analysis reveals that the primary crystalline phases include MgO, Mg₂SiO₄, Ca₃(PO₄)₂, and CaCO₃. During the early stages of the MAO process, the electrolysis of water in the alkaline electrolyte generates oxygen and hydrogen gases at the anode and cathode, respectively, leading to the formation of bubbles on the electrode surfaces. As the applied voltage increases, micro-arc discharges initiate and gradually break through the passivation coating on the alloy surface. Under the influence of strong localized electric fields, electrolyte solutes, particularly Ca²⁺ and Mg²⁺, interact with the molten substrate material within the discharge channels. This results in the formation of initial quantities of Ca₃(PO₄)₂ and Mg₂SiO₄. As the reaction progresses into the constant-current phase, elevated voltages intensify the plasma discharges, leading to ejection of molten substrate material from the discharge channels. This molten ejecta reacts with silicate (SiO₃²⁻) and phosphate (PO₄³⁻) ions in the electrolyte, promoting the formation and deposition of larger quantities of Mg₂SiO₄ and Ca₃(PO₄)₂ on the coating surface. The formation of CaCO₃ is attributed to reactions between Ca²⁺ and dissolved CO₂ from the electrolyte environment.

Despite the presence of Ag in the alloy substrate, no Ag-containing crystalline phases were detected in the XRD spectra, likely due to the low concentration of silver in the coating and its possible amorphous or dispersed elemental form. To further investigate the incorporation of Ag, XPS analysis was performed on the MAO coatings of ZQ alloys. The Ag 3d spectra confirm the presence of elemental Ag, with its concentration increasing in proportion to the Ag content in the alloy substrate (Figure 8b). This suggests successful incorporation of metallic Ag into the MAO coating, even if below the detection threshold of XRD. The Ca 2p spectra confirm the presence of Ca₃(PO₄)₂ and CaCO₃, in agreement with the XRD results (Figure 8c). Additionally, the P 2p spectra indicate the coexistence of metal phosphates and metal phosphides within the coating, further demonstrating the complex chemical environment created during the MAO process (Figure 8d).

3.2.3. Comprehensive Properties of ZQ 0.8-MAO Coating

Figure 9a illustrates the surface wear tracks and wear volumes of the ZQ 0.8-MAO coating, the data reveals that the having a smaller wear volume at just 1.11 × 10⁵ μm³. The ZQ 0.8-MAO coatings is denser and more uniform, with fewer micro-defects, resulting in better wear resistance. Additionally, the MAO process, which involves the synergistic effects of high voltage and high current, creates a ceramic coating that strongly bonds with the metal substrate, further reducing the wear volume when the friction pair contacts the dense MAO coating.

The adhesion strength testing of the coatings shown in Figure 9b supports this conclusion. In the figure, the parameter Lc represents the critical load at which the indenter penetrates through the MAO coating, while the AE curve denotes the acoustic emission signal curve. The macroscopic scratch trajectory is visible below the acoustic emission signal curve, with point (A) indicating the initial location where the coating is first breached. By analyzing and comparing the acoustic emission signal curve with the initial penetration location of the coating, it is evident that the acoustic emission signals of the four coatings exhibit significant fluctuations during the early stages of the test. This is due to the presence of heterogeneous hard phases and numerous defects within the MAO coatings, which cause unstable peeling of the coating during the load application, leading to pronounced fluctuations in the acoustic emission signals [44]. This phenomenon is also an inevitable result of the indenter passing through the porous coating of the MAO coating, which, to some extent, reflects the thickness of the porous coating.

As the test progresses, the acoustic emission signals gradually decrease and stabilize, indicating that the indenter has begun to reach the dense coating of the MAO coating. As seen in the figure, before the coating is fully penetrated, the ZQ 0.8-MAO coating exhibits a longer duration of stable acoustic emission signals, suggesting a significant advantage in the thickness of the dense coating, at last, this coating is withstanding a load of 10.37 N at the time of penetration. Thus, it can be concluded that the ZQ 0.8-MAO coating has a substantial advantage in terms of coating adhesion strength, which helps to prevent the coating from detaching from the substrate due to poor adhesion, thereby enhancing the effectiveness of the implant [45].

As a biomedical metal material, magnesium alloys implanted into the human body require not only certain biological properties but also sufficient corrosion resistance. This ensures that the magnesium alloy implants can function within the body for a certain period, providing adequate time for bone tissue to heal [46]. Figure 10 presents the electrochemical spectra and fitted circuits of the ZQ alloy MAO coatings, with

Table 3. Electrochemical fitting data of ZQ-MAO coating.

Sample	Rs (Ω·cm2)	CPE1 (Ω−1·cm−2 sn)	Rc (Ω·cm2)	Rct (Ω·cm2)	CPE2 (Ω−1·cm−2 sn)
ZQ 0.8-MAO	18.12	9.97 × 10−7	2.15 × 103	2.56 × 104	1.29 × 10−6

displaying the corresponding electrochemical impedance fitting data. In the fitted circuits, Rs represents the resistance of Hank’s solution, which remains within a certain range and does not significantly impact the corrosion resistance data of the coatings. R_c refers to the resistance of the coating in protecting the substrate, while R_ct represents the charge transfer resistance between the coating and the ZK60 substrate. The R_ct value is a reliable indicator of the overall corrosion resistance of the coating. Therefore, in this study, R_ct is considered as the standard for evaluating the corrosion resistance of the coatings. The data reveal that the coating prepared from the ZQ 0.8-MAO has a higher impedance value, reaching 2.56 × 10⁴ Ω·cm². In Figure 7b, the ZQ 0.8-MAO is thick, uniform, and has fewer surface and internal defects. Consequently, it exhibits the best corrosion resistance in electrochemical testing.

The primary cause of infections in biomaterials is the formation of bacterial biological coating [47]. These biofilms protect bacteria from the immune system and antibiotic treatments, leading to severe damage to implants [48]. Therefore, it is crucial for implants to possess some level of antibacterial capability to slow down or even prevent the formation of bacterial biofilms. Figure 11 illustrates the bacterial colony images from the antibacterial experiments and contact angle of MAO coatings on the ZQ 0.8 alloy. The analysis indicates that when a material’s water contact angle is less than 80°, it is considered to have good biocompatibility. The preparation of a coating on the alloy surfaces improves the wettability of the materials to some extent, as evidenced by the decrease in contact angles from approximately 100° before coating. Notably, the contact angle of the ZQ 0.8-MAO coating drops to 41.19°, indicating strong hydrophilicity, which is favorable for cell attachment and growth on the material’s surface. This demonstrates that the MAO coating enhances the wettability of the materials, thereby improving their biological performance.

Figure 11 clearly shows that the coating from the alloy containing 0.8 wt.% Ag can kill most of the bacteria. Only a few colonies remain on the plate, achieving an antibacterial rate of 96.2%, which meets the requirements for antibacterial performance in implants. Due to its potent bactericidal properties and broad-spectrum antibacterial characteristics, silver has garnered significant attention among various metal-based inorganic antibacterial agents. Research indicates that the antibacterial mechanisms of silver can be categorized into two main types [45]. Firstly, the Contact Reaction Mechanism, the cell membrane has a negative charge, which allows Ag⁺ ions to bind to it under the influence of Coulomb forces when they come into contact. Ag⁺ binds with proteins in the bacterial cell membrane, inducing the production of reactive oxygen species (ROS) [48], which are strong oxidizing agents with a high oxidative potential. This process damages the cell membrane, leading to the leakage of nutrients from the cell, thereby achieving a bactericidal effect. After the cell membrane is ruptured, Ag⁺ enters the cell and reacts with enzymes, coagulating enzyme proteins and disrupting the cell’s ability to synthesize essential nutrients. Ag⁺ also interferes with the formation of chitin, hindering cell wall formation and inhibiting cell reproduction and growth. Additionally, Ag⁺ can react with DNA inside the cell, directly blocking bacterial reproduction by damaging genetic factors. Crucially, after completing the bactericidal action, Ag⁺ can be released from the dead bacteria and repeat the antibacterial process, thereby establishing a prolonged antibacterial effect [46]. Secondly, the Photocatalytic Mechanism. Under light exposure, Ag nanoparticles and Ag⁺ can catalyze and activate water molecules and oxygen in the air, producing highly oxidative hydroxyl radicals and reactive oxygen ions. These substances can quickly disrupt the bacterial reproduction system, preventing bacteria from multiplying, thereby achieving an antibacterial effect.

In this experiment, a certain amount of Ag⁺ is present on the surface of the MAO coating. When the coating is cultured with bacterial suspension, the Ag⁺ on the surface is released into the suspension as the coating degrades, initiating a series of destructive actions against the bacteria. However, because the amount of Ag⁺ released is limited, the alloys with different Ag content exhibit varying degrees of antibacterial performance in the experiment.

4. Conclusions

The paper first prepared four novel ZQ alloys and assessed their mechanical properties, microstructure, corrosion resistance, and biocompatibility to evaluate their potential as biomaterials. Subsequently, MAO treatment was applied to the surface of the ZQ 0.8 alloy to enhance its corrosion resistance and biocompatibility. The main conclusions are as follows:

1.

The newly developed magnesium alloys have an elastic modulus closer to that of human bone, approximately 25.8 GPa, which can effectively reduce the stress shielding effect when implanted in the human body. The primary phases present in the alloy include the α-Mg matrix, Mg₂Ca, Mg₇Zn₃, and Mg₆Ca₂Zn₃ phases, with Mg₂Ca mainly distributed in the form of dots within the magnesium alloy matrix and Mg₆Ca₂Zn₃ distributed along the grain boundaries. The addition of alloying elements improved the mechanical properties of the magnesium alloys, making them more akin to human bone.

2.

Variations in phase content resulted in differences in corrosion resistance among the alloys. The ZQ 0.8 magnesium alloy exhibited the highest electrochemical impedance value of 57.52 Ω·cm², while the ZQ 0.2 magnesium alloy had a much lower value of 9.68 Ω·cm². The addition of Ag to the magnesium alloy conferred excellent antibacterial properties, with antibacterial rates exceeding 97%, meeting the requirements for biomedical materials.

3.

ZQ 0.8-MAO coating was compact with a smooth surface, low porosity, and substantial thickness, providing the best surface and cross-sectional morphology. Phase analysis revealed the presence of MgO, Mg₂SiO₄, Ca₃(PO₄)₂, CaCO₃, and Ag₂O on the surface of MAO coating. This MAO coating also exhibited the best adhesion to the substrate and excellent antibacterial properties, with an antibacterial rate of up to 96.2%. Electrochemical tests also revealed that ZQ 0.8-MAO coating demonstrated good corrosion resistance, with a fitted impedance value of 2.56 × 10⁴ Ω·cm².