Improving the Corrosion Resistance and Blood Compatibility of Magnesium Alloy via Fe-Based Amorphous Composite Coating Prepared by Magnetron Sputtering

Abstract

Magnesium alloy represents a typical category of biodegradable medical materials. However, the poor corrosion resistance and rapid degradation have significantly hindered the clinical adoption of magnesium alloy implants. This paper puts forward a method to improve the corrosion resistance of magnesium alloy by using an Fe-based composite coating. The microstructure and composition of the coating were analyzed by scanning electron microscopy (SEM), transmission electron microscopy (TEM), and energy-dispersive spectroscopy (EDS). The corrosion resistance was evaluated through potentiodynamic polarization and electrochemical impedance spectroscopy (EIS) measurements conducted in simulated body fluid, while the degradation behavior of the samples was evaluated by examining the hydrogen evolution volume and corrosion morphology during immersion tests. The results indicate that the composite coating exhibits a dual-layer structure, consisting of an amorphous carbon–fluorine transition layer and an iron-rich surface layer. After coating treatment, the corrosion current density of magnesium alloy decreased from 1.38 × 10⁻⁴ to 3.41 × 10⁻⁶ A/cm². Throughout a 28-day immersion period, the composite-coated sample demonstrated a remarkably low hydrogen evolution rate and maintained a smooth, intact surface. Furthermore, hemolysis and platelet adhesion tests confirmed the outstanding blood compatibility of the composite-coated magnesium alloy, showing an ultralow hemolysis rate of 0.1% and minimal platelet adhesion with well-preserved morphology.

1. Introduction

With the advancement of modern medicine, biodegradable biomedical materials have attracted widespread attention. As an emerging biodegradable metal for biomedical applications, magnesium alloys offer excellent mechanical properties, good biocompatibility, and cost-effectiveness, demonstrating broad prospects. However, the excessively rapid degradation rate of magnesium-based biomedical materials leads to a sharp decline in the mechanical integrity of implants [1,2,3,4]. Furthermore, the rapid degradation of magnesium implants can cause a localized increase in pH in body fluids accompanied by substantial hydrogen gas evolution, which may adversely affect surrounding tissue physiology [5]. It is evident that the high corrosion rate severely limits the clinical application of magnesium alloy biomedical materials.

Currently, optimizing the composition and microstructure of magnesium alloys by adding alloying elements or applying heat treatment can enhance the corrosion resistance [6,7,8]. Rare earth elements such as yttrium (Y) and neodymium (Nd) can refine grains, reduce grain boundary corrosion, and promote the formation of a dense oxide layer on the surface. For instance, WE43 magnesium alloy (Mg-4%Y-3%RE) exhibits significantly better corrosion resistance than pure magnesium and conventional magnesium alloys [9,10]. A recent study on additively manufactured WE43 alloys has proposed that high-temperature oxidation can reduce the corrosion rate of such alloys in simulated body fluid (SBF) [11].

However, most medical magnesium alloys cannot fully meet the requirements of biomedical applications without surface treatment. Surface treatment is one of the most direct and effective methods to improve the corrosion resistance and biocompatibility of magnesium alloy implants [12,13]. As a surface treatment approach, coating technology can form an effective protective layer on medical magnesium alloys, effectively blocking direct contact between the magnesium substrate and corrosive media. Currently, surface coatings applied to medical magnesium alloys or Mg-bearing implants are generally categorized into organic and inorganic coatings. Representative organic coatings include hyaluronic acid, polylactic acid (PLA), and poly(lactic-co-glycolic acid) (PLGA), while common inorganic coatings encompass fluorides, hydroxyapatite (HA), and titanium oxide, among others. These coatings significantly improve the corrosion resistance of magnesium alloys. Furthermore, certain types can provide excellent biocompatibility and even exhibit specific biological functions, such as antibacterial properties, drug delivery capability, and the promotion of tissue regeneration [14,15,16,17,18,19].

Fe-based biomedical materials, which take iron as their main component, have garnered increasing research attention owing to their favorable mechanical properties, good biocompatibility, and a slow degradation rate. Iron has a high elastic modulus, yield strength, and tensile strength, thereby meeting the mechanical requirements for implants applications [20,21,22]. Iron is an essential trace element in the human body, involved in various critical biochemical processes such as oxygen transport, DNA synthesis, cellular respiration, and hormone production. It plays a vital role in maintaining normal hematopoietic and immune functions, as well as promoting metabolism [23,24]. Iron degrades slowly into Fe²⁺ and Fe³⁺ ions in biological media without producing hydrogen gas during the slow degradation process [25,26,27]. Moreover, the hemocompatibility of iron and its alloys was favorable, with a hemolysis percentage below 5%, and the Fe-based coating exhibited enhanced blood compatibility over stainless steel [28,29,30,31]. Therefore, the Fe-based coating holds promising potential for in vivo applications.

Magnetron sputtering technology is characterized by high film density, minimal substrate temperature elevation, and ease of scale-up for production. These features make it well-suited for fabricating functional coatings. Therefore, in this study, a uniform and dense multifunctional Fe-based coating was prepared by magnetron sputtering on the surface of medical magnesium alloys to enhance their corrosion resistance and biocompatibility. The composition, morphology, corrosion resistance, and blood compatibility of the Fe-based coating were investigated, with the aim of effectively mitigating the drawbacks arising from the excessively rapid and uncontrolled degradation of medical magnesium alloy implants.

2. Materials and Methods

2.1. Preparation of the Coating

For this study, AZ31 magnesium alloy was selected, and its chemical composition is detailed in

Table 1. Chemical composition of AZ31 magnesium alloy (wt. %).

Al	Zn	Mn	Si	Fe	Ni	Cu	Mg
3.120	0.930	0.300	0.018	0.003	0.001	0.001	balance

. As-extruded AZ31 rods with a diameter of Φ 10 mm were cut into samples with a thickness of 3 mm. These samples were progressively ground using 100#, 200#, 400#, 800#, and 1000# SiC papers to remove surface oxides and defects, followed by fine polishing with 1 μm diamond paste to achieve a smooth surface. Finally, the polished samples were ultrasonically cleaned in anhydrous ethanol for 5 min to eliminate residual abrasive particles and polishing compounds.

Fe-based composite coating was deposited on the magnesium alloy substrate via magnetron sputtering at room temperature. A relatively low temperature can avoid the impact of high temperatures on the microstructure and properties of magnesium alloy. To mitigate potential galvanic corrosion between the iron coating and the Mg alloy substrate, a carbon–fluorine transition layer was first deposited by sputtering a high-purity polytetrafluoroethylene (PTFE) target using a radio-frequency magnetron sputtering system. Subsequently, the PTFE target was replaced with a 316L stainless steel target (purity: 99.95, composition listed in

Table 2. Chemical composition of 316L stainless steel target (wt. %).

Cr	Ni	Mo	Mn	Si	C	P	Fe
16–18	10–14	2–3	0.2–2	<1	<0.03	<0.045	balance

) to deposit the iron-based surface coating. The detailed magnetron sputtering parameters are summarized in

Table 3. Magnetron sputtering parameters of Fe-based composite coating.

Target	Temperature (°C)	Power (W)	Pressure (Pa)	Argon Flow (sccm)	Time
PTFE	25	120	0.6	30	5 h
316L	25	120	0.6	30	20 min

. Under the same conditions, 20 samples were fixed at the same horizontal height on a rotating round stage for each deposition run, ensuring a uniform coating deposition on each sample’s surface. The stage was set to rotate at 20 rpm. This experimental setup enables the measurement of coating thickness using only one representative sample.

2.2. Characterization

After the coating preparation, a sharp blade was used to carefully scratch the surface of the coating to expose the cross-section, and then the surface morphology and cross-sectional structure of the coating were observed using scanning electron microscopy (SEM). Under identical sputtering conditions, PTFE targets and 316L stainless steel targets were sputter-deposited onto TEM-grade copper grids. The microstructure and crystal structure of the coatings were characterized by transmission electron microscopy (TEM), and their elemental composition was analyzed via energy-dispersive spectroscopy (EDS).

2.3. Corrosion Resistance Experiments

Kokubo’s simulated body fluid (SBF) was employed as the corrosive medium [32]. The corrosion resistance of the samples was evaluated using an electrochemical workstation equipped with a standard three-electrode system, including a platinum electrode served as the counter electrode, the sample acted as the working electrode (with an exposed surface area of 1 cm² in SBF), and a saturated calomel electrode (SCE) functioned as the reference electrode. The electrochemical testing procedure began with monitoring the open-circuit potential (OCP) for 1 h to achieve a stable state. This was followed by electrochemical impedance spectroscopy (EIS) measurements, performed over a frequency range from 0.01 Hz to 100 kHz with an AC amplitude of 5 mV. Subsequently, potentiodynamic polarization tests were conducted at a scan rate of 0.5 mV/s [11]. For each test, three parallel specimens were measured, and the average value was taken.

For immersion testing, samples were immersed in SBF at a volume-to-surface-area ratio of 40 mL/cm² and maintained at 37 °C for 28 days. Following immersion, the samples were rinsed thoroughly with deionized water, dried in atmosphere, and then examined using SEM and EDS to evaluate the corrosion morphology and elemental composition. Throughout the immersion period, the hydrogen gas was continuously collected to evaluate the degradation rate of the samples.

2.4. Blood Compatibility Evaluation

Commercially available rabbit blood was used for the hemolysis rate test. First, 0.109 M sodium citrate anticoagulant and blood were mixed at a volume ratio of 1:9 for anticoagulation treatment. Then, the anticoagulated blood was diluted with physiological saline at a volume ratio of 4:5. The ultraviolet-sterilized samples (n = 5) were placed in polypropylene centrifuge tubes, and 10 mL of physiological saline was added to each tube. Two additional centrifuge tubes were prepared separately: one with an equal volume of physiological saline (negative control) and the other with an equal volume of distilled water (positive control). All tubes were preheated at 37 °C for 30 min, after which 0.2 mL of diluted blood was added to each tube and mixed thoroughly. The tubes were then maintained at 37 °C for another 60 min. After that, the tubes were taken out for centrifugation. The supernatant was aspirated, and the absorbance value was measured at a wavelength of 545 nm using an ultraviolet spectrophotometer. The hemolysis rate of the samples was calculated according to Equation (1):

$$HR(\%)={\displaystyle \frac{D_t-D_{nc}}{D_{pc}-D_{nc}}}\times 100\%$$

(1)

where HR refers to the hemolysis rate; D_t, D_nc, and D_pc represent the absorbance values of the experimental group, negative control group, and positive control group, respectively.

The sterilized samples were placed in a 24-well cell culture plate. Next, 500 μL of platelet-rich plasma (PRP) was aspirated and evenly dropped onto the sample surfaces. After incubation at 37 °C for 120 min, the samples were gently rinsed with physiological saline 3–5 times to remove non-adherent platelets. Subsequently, the samples were immersed in a 0.2% glutaraldehyde solution to fix the surface-adherent platelets for 6 h. Gradient dehydration was then performed sequentially. Once the samples were completely dried, the adhesion of platelets on their surfaces was observed using a scanning electron microscope.

3. Results and Discussion

3.1. Characteristics of the Coating

To clarify the microstructure of the transition layer obtained by sputtering the PTFE target, PTFE was sputter-deposited onto a TEM copper grid under the same preparation conditions, and the deposited layer was characterized using TEM. The high-resolution TEM (HRTEM) image (Figure 1a) shows that no lattice fringes with a specific spacing and orientation were observed in the deposited layer. In addition, the selected area electron diffraction (SAED) pattern of the deposited layer exhibits a typical amorphous diffraction characteristic, i.e., a continuously distributed halo ring (Figure 1b), confirming that the deposited layer has an amorphous structure. The EDS results (Figure 1c) show that the atomic percentage of C and F elements in the deposited layer is close to the stoichiometric ratio of 1:2, which is consistent with the elemental ratio in the PTFE target. Elemental mapping analysis confirms that C and F elements are highly uniformly distributed in the coating, indicating that an amorphous C-F layer with uniform composition was obtained by magnetron sputtering.

To obtain the microstructural characteristics of the Fe-based coating, 316L target was sputter-deposited onto a copper grid under the same process conditions. HRTEM analysis (Figure 2a) shows that there are no periodic lattice fringes in the Fe-based coating, and only nanoclusters with a size of less than 2 nm are observed. The SAED pattern exhibits the characteristic of a typical amorphous continuous diffuse ring, with no diffraction spots of crystals detected (Figure 2b). Due to the severe inhibition of atomic diffusion ability under room-temperature deposition conditions, the prepared Fe-based coating possesses typical amorphous structural features. As shown in Figure 2c, EDS results show that the atomic percentages of the three main elements (Fe, Cr, and Ni) in the coating are 54.5%, 11.6%, and 6.8%, respectively, which are close to the composition of the 316L stainless steel target. Elemental mapping analysis indicates that all elements are highly uniformly distributed in the coating. Studies have demonstrated that iron-based amorphous coatings possess high corrosion resistance, which is primarily attributed to the absence of crystalline defects—including dislocations and grain boundaries—in their microstructures [33,34].

The surface and cross-section of the Fe-based composite coating were analyzed using SEM, as shown in Figure 3. The deposition times of the C-F intermediate layer and the Fe-based coating were 5 h and 20 min, respectively. The result indicates that the composite coating exhibits a typical dual-layer structure: the intermediate layer is the amorphous C-F layer with a thickness of approximately 700 nm, and the surface layer is an amorphous Fe-based coating with a thickness of approximately 100 nm, resulting in a total thickness of the composite coating of about 800 nm. The deposition rate of Fe-based coating is about 5 nm/min. It was found that the boundary between the intermediate layer and the surface layer are clear and flat, and the bonded interfaces fit tightly together without any signs of detachment. The microstructure of the Fe-based coating is granular, comprising nanoscale particles. This morphology results in exceptional surface flatness and structural density. The inherent nature of magnetron sputtering dictates that the deposited coating exhibits excellent coverage and adhesion to the substrate. Specifically, when high-energy argon Ar⁺ ions bombard the target surface under the influence of an electric field, atoms in the target are ejected from its surface and deposited onto the substrate. As a large number of these deposited atoms agglomerate, a continuous coating is formed on the substrate surface. This dense coating can effectively inhibit the penetration of corrosive media toward the substrate, thereby exerting a positive effect on the corrosion protection of the magnesium alloy substrate. Additionally, the amorphous structure also contributes positively to enhancing the corrosion resistance of the iron-based coating itself, as analyzed earlier. This, in turn, enables a further enhancement of the protective efficacy of the thin iron-based coating on the substrate.

3.2. Corrosion Behavior

As shown in Figure 4, the corrosion behavior of AZ31 magnesium alloys with and without Fe-based composite coatings was quantitatively evaluated using electrochemical potentiodynamic polarization curves. Corrosion potential (E_corr) and corrosion current density (I_corr) can be derived from the polarization curves. Generally, a higher E_corr and a lower I_corr indicate better corrosion resistance and a slower corrosion rate of the material [35]. The results show that after depositing the Fe-based composite coating, the E_corr of the AZ31 sample increased from −1.50 V to −1.46 V. The coated AZ31 exhibited a more positive corrosion potential compared to the bare AZ31 substrate, suggesting that the bare AZ31 magnesium alloy is more susceptible to corrosion than the sample with Fe-based composite coating. Additionally, the I_corr of the sample coated with Fe-based composite coatings was significantly lower than that of the bare AZ31 sample, decreasing from 1.38 × 10⁻⁴ A/cm² to 3.41 × 10⁻⁶ A/cm². Therefore, the Fe-based composite coating developed in this study is an excellent protective coating that can effectively reduce the corrosion rate of the AZ31 substrate.

Figure 5 shows the Nyquist plots of electrochemical impedance spectroscopy for the samples after immersion in simulated body fluid for 1 h. As observed from the shape of the Nyquist plots, the curve of bare AZ31 magnesium alloy consists of a capacitive arc in the high-frequency region and a capacitive arc in the medium-frequency region. However, the coated AZ31 only have a single capacitive loop. Studies have indicated that the high-frequency capacitive arc is related to the electron transfer process of the electric double layer on the magnesium alloy surface, while the medium-frequency capacitive arc is associated with the electrode behavior of the passive film on the substrate surface. To quantitatively analyze the electrochemical corrosion properties of the samples, Zview2 software was used to fit the EIS data of the samples. In the fitting model, R_s is the solution resistance, R₁ represents the corrosion product film resistance on the sample surface, R_ct represents the charge transfer resistance, and CPE represents constant phase angle elements. The fitting parameters were obtained in

Table 4. Fitted results obtained from the equivalent circuit of electrochemical impedance spectra.

Sample	Bare AZ31	Coated AZ31
Rs/(Ω·cm2)	30.17	18.36
R1/(Ω·cm2)	63.31	-
CPE1-T/(F·cm2)	2.15 × 10−3	-
CPE1-P/(F·cm2)	1.023	-
Rct/(Ω·cm2)	141.40	134,770
CPE2-T/(F·cm2)	1.10 × 10−5	6.71 × 10−5
CPE2-P/(F·cm2)	0.86	0.95

. The fitting results revealed that the charge transfer resistance of the AZ31 substrate was 141.4 Ω·cm², whereas that of the coated AZ31 magnesium alloy reached as high as 1.34 × 10⁵ Ω·cm². Since R_ct reflects the resistance during charge transfer at the electrode/electrolyte interface, this indicates that the magnesium alloy substrate has poor corrosion resistance, while the Fe-based composite coating significantly improves the corrosion resistance of the AZ31 magnesium alloy.

To clarify the variation law of the corrosion rate of samples during immersion, the hydrogen evolution amount of AZ31 samples with and without Fe-based coating were measured. As indicated by the curves in Figure 6, the AZ31 magnesium alloy substrate continuously released hydrogen throughout the immersion period; in particular, the hydrogen evolution rate was relatively fast at the initial stage of immersion. Although the rate slowed down afterward, hydrogen release remained continuous, with a total hydrogen evolution amount of approximately 12.7 mL after 28 days. In sharp contrast, the AZ31 magnesium alloy with Fe-based composite coating exhibited an extremely low hydrogen evolution rate throughout the entire immersion process, and the total hydrogen evolution amount collected after 28 days was only about 0.3 mL. It can be concluded that the Fe-based composite coating has achieved remarkable effects in regulating the corrosion rate of magnesium alloy. First, it significantly reduces the overall corrosion rate of the magnesium alloy. Second, it effectively inhibits the excessive early-stage corrosion of the magnesium alloy, thereby stabilizing the initial degradation rate of the magnesium alloy substrate.

Figure 7a,b show the corrosion morphologies of the magnesium alloy substrate after 28 days of immersion in SBF. As observed in Figure 7a, due to the long-term accumulation of localized pitting corrosion, large and deep corrosion pits formed on the sample surface. Under such circumstances, if the magnesium alloy is used as an in vivo implant, its overall mechanical properties will be rapidly lost. Figure 7b shows the surface morphologies of the magnesium alloy with Fe-based composite coating after 28 days of immersion. It can be seen that no severe localized corrosion was found on the surface of the sample. Figure 7c shows the initial surface of the coating prior to corrosion, which serves as a direct reference for comparison with the corroded surface. At the same magnification scale as in Figure 7d, after 28 days of corrosion exposure, the Fe-based coating exhibited uniform cracking; however, it still maintained a sound bonding state with the substrate. As shown in Figure 7e,f, the uniform distribution of Fe element confirms that the composite coating has good macroscopic integrity. Meanwhile, the detection signal of F element is significantly enhanced at the gaps of the Fe-based coating, which indicates that the C–F interlayer remains intact. Based on the current results, it is speculated that the corrosion of the magnesium alloy will accelerate after approximately one month due to coating cracking. Therefore, in future research, it is necessary to evaluate the corrosion behavior of the coating during longer-term service. In fact, controlling the degradation of biomedical magnesium alloy implants is a complex engineering challenge. It requires the development of protective coatings tailored to the duration of the patient’s treatment cycle. For instance, different coating thicknesses can be designed to achieve varying degradation times.

From the above experimental results, it can be concluded that the corrosion resistance of the magnesium alloy substrate is significantly enhanced owing to the effective protection of the Fe-based coating. After the Fe-based coating is applied, the corrosion current density of the magnesium alloy substrate is reduced by nearly two orders of magnitude, and the charge transfer resistance exhibits the same trend. Meanwhile, in the immersion test, it was observed that the amount of hydrogen gas generated by magnesium alloy corrosion is extremely low, which is attributed to the protection of the coating. These phenomena confirm the effectiveness of the Fe-based coating in reducing the corrosion rate of the magnesium alloy. We believe the above results are closely related to the composition and morphology of the Fe-based coating. First, the amorphous structure of the coating confers a corrosion resistance advantage. Second, the coating exhibits excellent densification, which enables it to fully cover the substrate and prevent corrosive media from easily penetrating to corrode the substrate.

3.3. Blood Compatibility

As shown in

Table 5. The results of hemolysis of samples.

Samples	Hemolysis Rate (%)
Bare AZ31	42.2 ± 2.5
Fe-coated AZ31	0.1 ± 0.08

, the magnesium alloy substrate exhibited severe hemolysis with a hemolysis rate as high as 42.2%. In contrast, the Fe-based composite-coated magnesium alloy showed significantly improved hemocompatibility, with a hemolysis rate of only 0.1%, which meets the safety standards for medical implants. During the experiment, it was observed that when the magnesium alloy came into contact with physiological saline, rapid corrosion occurred. As the concentration of Mg²⁺ increased, a large amount of Mg²⁺ penetrated the red blood cell membrane and entered the cell interior, causing the intracellular ion concentration to exceed the normal range. To dilute the excessively high intracellular ion concentration, red blood cells regulated osmotic pressure by absorbing water, and eventually ruptured due to excessive swelling, thereby inducing hemolysis. Since the hemolysis rate of the magnesium alloy with Fe-based composite coating was significantly reduced, it can be seen that the Fe-based composite coating not only greatly reduces the corrosion rate of the magnesium alloy but also effectively improves the severe hemolysis problem of the magnesium alloy substrate.

Figure 8a shows the adhesion morphology of platelets on the surface of the AZ31 magnesium alloy substrate. As observed from the figure, the AZ31 alloy surface exhibited severe platelet adhesion following a 1 h culture period. This was evidenced by both a high density of adherent platelets and the presence of obvious aggregation. The adhesion and aggregation of platelets will promote their further activation, thereby inducing thrombosis. This indicates that the uncoated AZ31 magnesium alloy substrate is prone to platelet adsorption and exhibits poor hemocompatibility. Figure 8b presents the adhesion morphology of platelets on the surface of the Fe-based composite coating. Compared with the bare substrate, the number of platelets adhering to the composite coating was significantly reduced, with no aggregation observed. And the sparsely distributed platelets were nearly round in shape, which indicates that the platelets on the coating surface were not activated. In general, flat surfaces with nanoscale roughness exhibit an inhibitory effect on platelet adhesion [16]. The Fe-based coating fabricated in this study is characterized by nanoscale particles and possesses a highly flat, smooth surface. Furthermore, from the perspective of surface energy, the coating demonstrates a smaller contact angle with blood than the magnesium alloy substrate—an observation indicative of favorable hydrophilicity—which also contributes to inhibiting platelet adhesion. In conclusion, the Fe-based composite coating possesses superior anti-platelet adhesion capability and demonstrates good application potential in the field of medical implants.

4. Conclusions

In this study, an Fe-based composite coating was deposited on the surface of AZ31 magnesium alloy via magnetron sputtering. The structure, corrosion behavior, and blood compatibility of the coating were investigated. The key conclusions drawn from this work are summarized as follows:

• After 20 min of deposition at room temperature, a continuous, dense Fe-based coating with an amorphous structure was obtained, with a thickness of approximately 100 nanometers. The microstructure of the Fe-based coating is granular, comprising nanoscale particles.
• After the Fe-based coating treatment, the corrosion resistance of the AZ31 magnesium alloy substrate is significantly improved. The corrosion current density decreases from 1.38 × 10⁻⁴ A/cm² to 3.41 × 10⁻⁶ A/cm² in SBF and the charge transfer resistance increases remarkably. Compared with the uncoated substrate, the coated AZ31 magnesium alloy exhibits an extremely low hydrogen evolution rate during the 28-day immersion test, and its surface morphology remains intact after corrosion.
• The uncoated magnesium alloy induced severe hemolysis and platelet aggregation, whereas the Fe-based composite coating exhibited a hemolysis rate of merely 0.1% and superior anti-platelet adhesion capability—indicating excellent hemocompatibility.

Therefore, the Fe-based coating deposited on magnesium alloy via magnetron sputtering in this study holds further promising potential for in vivo applications.