A magnesium-based medical material containing a Zn / ZnO@HA / PCL biological coating and a preparation method and application thereof

Abstract

This invention discloses a magnesium-based medical material containing a Zn / ZnO@HA / PCL bio-coating, its preparation method, and its application, belonging to the technical fields of metal surface engineering and biomedical materials. The magnesium alloy-based medical material comprises a magnesium alloy substrate and a Zn transition layer, a ZnO@HA layer, and a PCL polymer layer sequentially stacked on the magnesium alloy substrate. ZnO@HA is a hydroxyapatite-zinc oxide composite material. This composite coating material, through the synergistic construction of the zinc transition layer, the inorganic functional ZnO@HA layer, and the polymer PCL toughening layer, comprehensively enhances the corrosion resistance, antibacterial properties, and bioactivity of the magnesium alloy surface.

Description

Technical Field

[0001] This invention relates to a magnesium-based medical material containing a Zn / ZnO@HA / PCL bio-coating, its preparation method and application, belonging to the technical fields of metal surface engineering and biomedical materials. Background Technology

[0002] Magnesium alloys possess advantages such as low density, high specific strength, good biocompatibility, and biodegradability, making them promising candidates for orthopedic implant materials. Particularly in biodegradable bone fixation devices, magnesium alloys can gradually degrade and be absorbed by the body, thus avoiding the risk of secondary surgery to remove the implant. However, magnesium alloys exhibit high chemical reactivity in physiological environments, resulting in rapid corrosion and quick degradation. This leads to premature failure of mechanical properties, releases large amounts of hydrogen gas, and causes local alkalization, affecting the normal healing of surrounding tissues. Therefore, effectively controlling the corrosion behavior of magnesium alloys is a key issue in their clinical application.

[0003] Currently, surface modification technology is considered an effective way to improve the corrosion resistance of magnesium alloys. By constructing functional coatings on the surface of magnesium alloys, it is possible not only to effectively isolate external corrosive media and slow down the degradation rate of the matrix, but also to endow the material with more functional properties, such as biocompatibility, antibacterial properties, and osteogenic activity. Among them, hydroxyapatite (HA) is widely used as a surface coating for biomedical magnesium alloys because its chemical composition is highly similar to the mineral phase of natural bone tissue and it has excellent bioactivity and osteoconductivity. However, single HA coatings still face many challenges in practical applications: First, HA is inherently brittle and is prone to cracking or even peeling under load or deformation conditions; second, HA itself lacks antibacterial function and is difficult to inhibit the risk of postoperative bacterial infection; third, the bonding strength between the HA coating and the magnesium alloy matrix is ​​limited, and the coating is prone to premature detachment due to interfacial failure in the body fluid environment.

[0004] Therefore, developing a composite coating system with stable structure, antibacterial ability, bioactivity, and good corrosion resistance is of great significance for improving the application value of magnesium alloys in the biomedical field. Summary of the Invention

[0005] In view of the problems existing in the prior art, one of the objectives of the present invention is to provide a magnesium-based medical material containing a Zn / ZnO@HA / PCL bio-coating. This composite coating material achieves a comprehensive improvement in the corrosion resistance, antibacterial properties and bioactivity of the magnesium alloy surface through the synergistic construction of a zinc transition layer, an inorganic functional ZnO@HA layer and a polymer PCL toughening layer.

[0006] The second objective of this invention is to provide a method for preparing magnesium-based medical materials containing a ZnO@HA / PCL bio-coating, which is simple and reproducible.

[0007] A third objective of this invention is to provide a magnesium-based medical material containing a ZnO@HA / PCL bio-coating for use as an orthopedic implant. This material not only significantly improves the bonding strength between the HA coating and the matrix, but also effectively overcomes the inherent brittleness of bioceramic coatings. Simultaneously, it endows the material with excellent antibacterial properties and biocompatibility, thereby improving the corrosion resistance of magnesium alloys while ensuring the material's mechanical support capacity during bone tissue repair, meeting the modern medical demand for multifunctional orthopedic implant materials.

[0008] To achieve the above objectives, a first aspect of the present invention is to provide a magnesium-based medical material containing a Zn / ZnO@HA / PCL bio-coating, wherein the magnesium alloy-based medical material comprises a magnesium alloy substrate and a Zn transition layer, a ZnO@HA layer and a PCL polymer layer sequentially stacked on the magnesium alloy substrate; ZnO@HA is a hydroxyapatite-zinc oxide composite material.

[0009] The innovation of this invention lies in two aspects: First, it enhances the interfacial bonding between the coating and the substrate through a Zn transition layer. A Zn metal transition layer is first deposited on the surface of the magnesium alloy substrate. As a metallic material, Zn can form a good metallic bonding interface with the magnesium alloy, thereby effectively improving the adhesion between the coating and the substrate. Simultaneously, the Zn transition layer can improve the chemical activity and surface condition of the magnesium alloy surface, providing stable and uniform substrate conditions for the subsequent deposition of functional coatings, thus improving the structural stability and interfacial bonding strength of the entire composite coating system. Specifically, the Zn layer can block corrosive media (such as water, Cl-). - H + Direct contact between the Zn layer and the magnesium substrate can alter the electrochemical reaction pathway on the magnesium alloy surface and reduce rapid hydrogen evolution, thereby significantly improving the corrosion resistance and degradation controllability of the magnesium alloy. The Zn layer can buffer stress, reduce crack propagation, and improve structural stability, thus enhancing the overall coating durability. In this case, the Zn layer can also serve as an excellent transition layer, not only improving the coating bonding strength but also providing a favorable growth environment for the subsequent HA coating.

[0010] Secondly, a ZnO@HA composite structure is constructed through a hydrothermal reaction. During the hydrothermal reaction, ZnO structures are partially generated in situ on the surface of the Zn layer under high temperature and pressure. Simultaneously, HA undergoes heterogeneous nucleation on the Zn and ZnO surfaces and gradually grows. ZnO and HA combine during deposition to form a ZnO@HA functional coating with composite (micro / nano) structure characteristics. This structure can simultaneously perform multiple functions. The HA layer provides excellent bioactivity; HA is a bioceramic material highly similar to human bone tissue, exhibiting excellent biocompatibility and osteoinductive properties. By constructing an HA layer on the magnesium alloy surface, the bioactivity of the material surface can be effectively improved, thereby promoting cell adhesion, proliferation, and bone tissue regeneration, making the material more suitable for biomedical applications. The Zn layer and ZnO structure endow the coating with antibacterial properties, and the Zn layer and the ZnO structure generated in situ during the hydrothermal reaction can release Zn to a certain extent. 2+ These ions can interfere with bacterial cell membrane structure and metabolic processes, thereby inhibiting various bacteria. Therefore, the introduction of the ZnO structure not only enhances the functionality of the coating but also reduces the risk of infection in implanted materials to some extent.

[0011] Third, the brittleness of the ceramic coating can be improved by using a flexible PCL layer. Since HA is a ceramic material, its inherent brittleness may cause cracks or even localized damage to the coating when subjected to external stress. By introducing a biodegradable PCL polymer layer onto the surface of the ZnO@HA composite coating, a protective film with good flexibility can be formed, thereby alleviating the brittleness of the ceramic layer and effectively inhibiting the initiation and propagation of cracks.

[0012] This invention constructs a multi-layer composite structure consisting of a Zn transition layer, a ZnO@HA functional layer, and a PCL flexible protective layer, enabling different materials to complement and reinforce each other functionally, thereby forming a multi-functional composite coating system and achieving systematic optimization of the surface properties of magnesium alloys.

[0013] In this invention, it should be noted that the Zn transition layer coating is attached to the magnesium alloy substrate, the ZnO@HA layer is attached to the Zn transition layer, and the PCL polymer layer is attached to the ZnO@HA layer.

[0014] As a preferred embodiment, the thickness of the Zn transition layer is 1~10 µm.

[0015] As a preferred embodiment, the thickness of the ZnO@HA layer is 20~50 µm.

[0016] As a preferred embodiment, the thickness of the PCL polymer is 5~10 µm.

[0017] As a preferred option, ZnO and / or HA are attached to the Zn transition layer.

[0018] In this invention, HA refers to hydroxyapatite, and PCL refers to polycaprolactone.

[0019] As a preferred embodiment, the magnesium alloy substrate is selected from at least one of WE43 magnesium alloy, AZ31B magnesium alloy, ZK60 magnesium alloy, and pure magnesium.

[0020] As a more preferred embodiment, the magnesium alloy substrate is AZ31B magnesium alloy.

[0021] As a preferred embodiment, the coating adhesion of the magnesium-based medical material containing ZnO@HA / PCL bio-coating is 11~14 N, and the plasticity index is 0.5~0.8.

[0022] As a preferred embodiment, the corrosion potential E of the magnesium-based medical material containing the ZnO@HA / PCL bio-coating is... corr The corrosion current density is -0.45 to -0.25 V. corr It is 2.24 × 10 -10 ~4.50 × 10 -9 A / cm 2 .

[0023] A second aspect of the present invention is to provide a method for preparing a magnesium-based medical material containing a ZnO@HA / PCL bio-coating as described in the first aspect, comprising forming a Zn transition layer sequentially on a magnesium alloy substrate, followed by forming a ZnO@HA layer and a PCL polymer layer.

[0024] As a preferred embodiment, the Zn transition layer is prepared by vapor deposition.

[0025] As a preferred embodiment, the ZnO@HA layer is prepared by hydrothermal synthesis; and the ZnO is generated in situ from the Zn transition layer during the hydrothermal reaction; the HA is formed by heterogeneous nucleation and growth on the Zn transition layer or the ZnO surface.

[0026] As a preferred embodiment, the PCL polymer layer is prepared by spin coating. PCL material possesses excellent flexibility and film-forming properties, enabling the formation of a continuous and dense flexible protective layer on the coating surface. This effectively alleviates the brittleness of the ceramic layer, improves the crack resistance and structural stability of the coating, and thus results in a more comprehensive Zn / ZnO@HA / PCL composite coating exhibiting superior overall performance in terms of corrosion resistance, bioactivity, antibacterial properties, and structural stability.

[0027] As a preferred embodiment, the preparation method of the magnesium-based medical material containing ZnO@HA / PCL bio-coating includes the following steps:

[0028] (1) A Zn transition layer is formed on a magnesium alloy substrate by physical vapor deposition to obtain a magnesium alloy containing a Zn transition layer;

[0029] (2) A magnesium alloy containing a Zn transition layer is placed in a hydrothermal solution and subjected to hydrothermal reaction to deposit a magnesium alloy containing Zn / ZnO@HA.

[0030] (3) Polycaprolactone colloid is coated onto the surface of a magnesium alloy coating containing Zn / ZnO@HA by spin coating and cured.

[0031] As a preferred embodiment, the magnesium alloy substrate is sputtered before the physical vapor deposition method.

[0032] As a preferred embodiment, the magnesium alloy substrate is pretreated before sputtering. The surface pretreatment method includes grinding and polishing the magnesium alloy substrate with sandpaper, followed by ultrasonic cleaning.

[0033] It should be noted that the present invention does not have special requirements for the medium of the polishing fluid, and anhydrous methanol is preferred.

[0034] As a preferred embodiment, the negative bias voltage of the sputtering is sequentially adjusted to -800 ± 50 V, -600 ± 50 V, -400 ± 50 V, and -200 ± 50 V, with the deposition time for each negative bias voltage point being independently 20~40 s. More preferably, the deposition time for each negative bias voltage point is 40 s.

[0035] As a preferred embodiment, the physical vapor deposition method is selected from one of magnetically filtered cathode vacuum arc deposition, multi-arc ion plating, and magnetron sputtering.

[0036] As a more preferred embodiment, the physical vapor deposition method is magnetically filtered cathode vacuum arc deposition.

[0037] As a preferred embodiment, the process of the magnetically filtered cathode vacuum arc deposition method is as follows: controlling the vacuum level of the vacuum chamber of the magnetically filtered cathode vacuum arc deposition equipment to be less than or equal to 5 × 10⁻⁶. -3 After Pa, using a Zn target as an arc source, during the vapor deposition process, the arc initiation current was controlled at 60~120 A, the bending magnetic field current at 1.5~3.0 A, the beam current intensity at 350~400 mA, the negative bias voltage at 90 V~100 V, the duty cycle at 60~90%, the deposition time at 10~30 min, and the vacuum environment cooling time at >1 h.

[0038] As a more preferred embodiment, the process of the magnetically filtered cathode vacuum arc deposition method is as follows: controlling the vacuum level of the vacuum chamber of the magnetically filtered cathode vacuum arc deposition equipment to be less than or equal to 4 × 10⁻⁶. -3 After Pa, using a Zn target as the arc source, during the vapor deposition process, the arc initiation current was controlled at 85–95 A, the bent tube magnetic field current at 1.8–2.2 A, the straight tube magnetic field current at 3–5 A, the beam current intensity at 400 mA, the duty cycle at 90%, the deposition time at 15–25 min, and the vacuum cooling time at 12–24 h. More preferably, the arc initiation current was 90 A, the bent tube magnetic field current was 2 A, the deposition time was 15 min, and the vacuum cooling time was 24 h.

[0039] As a preferred embodiment, the purity of the Zn target material is not less than 95 wt%. A further preferred purity is ≥99.99 wt%.

[0040] As a preferred embodiment, the hydrothermal solution contains soluble salts containing calcium and soluble salts containing phosphate, and the atomic ratio of Ca to P is 1:1.66~1.68.

[0041] As a preferred embodiment, the calcium-containing soluble salt is calcium ethylenediaminetetraacetate, and the phosphate-containing soluble salt is potassium dihydrogen phosphate.

[0042] As a preferred embodiment, in the hydrothermal solution, the concentration of the calcium-containing soluble salt is 0.1~0.2 mol / L, and the concentration of the phosphate-containing soluble salt is 0.2~0.3 mol / L.

[0043] In this invention, Ca-EDTA refers to calcium ethylenediaminetetraacetate.

[0044] In this invention, the pH adjuster for adjusting the pH value of the hydrothermal solution is ammonia and / or sodium hydroxide solution.

[0045] As a preferred embodiment, the pH value of the hydrothermal solution is 8.5–11, the hydrothermal reaction temperature is 100–150℃, and the reaction time is 12–16 h. The pH value in the hydrothermal system is a crucial factor affecting HA crystal formation and Zn oxidation behavior, including HA crystal nucleation and deposition, and Zn to ZnO conversion. Hydrothermal temperature is a key factor controlling reaction kinetics and crystal growth rate; a suitable temperature can promote HA crystal growth and improve the crystallinity of the coating. Excessive temperature may lead to rapid crystal growth, a rough coating structure, and local structural inhomogeneity, thus affecting the quality of the HA coating. The hydrothermal reaction time primarily affects the coating thickness, crystal size, and structural integrity. In actual hydrothermal reactions, pH, temperature, and reaction time do not act independently but collectively influence the coating formation process. Therefore, by rationally controlling the pH value, reaction temperature, and reaction time of the hydrothermal reaction system, the structural morphology and deposition quality of the ZnO@HA composite coating can be effectively controlled, thereby obtaining a structurally uniform and stable composite functional coating.

[0046] As a more preferred embodiment, the hydrothermal reaction is carried out at a temperature of 120 °C for a duration of 15 h.

[0047] As a preferred embodiment, the number-average molecular weight of polycaprolactone in the polycaprolactone colloid is 40,000 to 60,000; the solvent in the polycaprolactone colloid is dichloromethane. Selecting PCL with a molecular weight of 40,000 to 60,000 achieves a good balance between film-forming properties, mechanical properties, processing rheology, and degradation behavior. On the one hand, medium molecular weight PCL has moderate chain entanglement and flowability, which is beneficial for forming a dense and uniform protective layer and effectively penetrating micro- and nano-structures to achieve pore sealing; on the other hand, its degradation rate is moderate, which can delay the initial corrosion of the matrix while ensuring the biodegradability of the material in the later stages. In contrast, low molecular weight PCL is prone to film-forming defects and rapid failure, while high molecular weight PCL has problems such as processing difficulties, increased interfacial stress, and slow degradation.

[0048] The present invention does not have special requirements for the preparation process of polycaprolactone colloid; any process known in the art can be used.

[0049] As a preferred embodiment, the concentration of the polycaprolactone colloid is 5-10 wt%, the spin coating rate is 500-1200 rpm, and the spin coating time is 30-100 s. The concentration of the polycaprolactone colloid is further preferably 7.5 wt%. The concentration of the polycaprolactone colloid, the spin coating rate, and the spin coating time all significantly affect the formation quality of the PCL coating. Specifically, the concentration of the polycaprolactone colloid primarily affects the film-forming ability and coating thickness. When the concentration of the polycaprolactone colloid is too low, it leads to uneven coating or insufficient local coverage; while when the concentration of the polycaprolactone colloid is too high, it may affect the uniformity of the coating and the stability of the surface structure. Furthermore, a higher spin coating rate enhances the centrifugal force, allowing the solution to spread evenly on the substrate surface, thus forming a more uniform and moderately thick PCL coating; while a lower spin coating rate may result in an excessively thick coating and localized accumulation. However, an excessively high spin coating rate can lead to enhanced centrifugal force, causing the solution to be excessively ejected from the substrate surface, resulting in a coating that is too thin or even partially exposed. Meanwhile, excessive shearing can disrupt the stable spreading of the solution on the surface, easily causing defects such as coating discontinuity or streaks. Furthermore, spin-coating time also affects the degree of solution spreading on the substrate surface and the solvent evaporation process. Appropriately extending the spin-coating time is beneficial for uniform solution distribution and the formation of a continuous and dense polymer film. Therefore, by rationally controlling the PCL solution concentration, spin-coating speed, and spin-coating time, a PCL coating layer with uniform structure, moderate thickness, and stable adhesion can be obtained, thereby further improving the overall structural stability and comprehensive performance of magnesium-based medical materials containing ZnO@HA / PCL bio-coatings.

[0050] As a preferred embodiment, the spin coating operation specifically includes: first, spin coating at 500-750 rpm for 30-80 seconds, and then spin coating at 800-1000 rpm for 20-60 seconds. The low-speed stage mainly promotes the full wetting and capillary penetration of the PCL solution into the porous structure of ZnO@HA, enabling it to penetrate deep into the micro-nano pores to achieve interface filling and physical sealing; subsequently, high-speed spin coating quickly removes excess solution and precisely controls the film thickness, thereby forming a uniform and dense polymer protective layer. The two-step spin coating process achieves synergistic control of "interfacial penetration - surface film formation," which is a key process step in obtaining a high-quality composite coating.

[0051] As a preferred method, before spin coating, the magnesium alloy containing Zn / ZnO@HA is immersed in a mixed solution of silane coupling agent and ethanol with a pH of 4-5.5 for 4-10 min.

[0052] As a preferred embodiment, the concentration of the silane coupling agent in the mixed solution is 15-25 g / L, and the concentration of ethanol is 85-95 wt%.

[0053] In this invention, the pH adjuster for adjusting the pH value of the mixed solution is glacial acetic acid.

[0054] The third aspect of this invention relates to the application of the magnesium-based medical material containing a ZnO@HA / PCL bio-coating as described in the first aspect or the magnesium-based medical material containing a ZnO@HA / PCL bio-coating prepared by the preparation method described in the second aspect as an orthopedic repair implant.

[0055] Compared with the prior art, the present invention has at least the following advantages:

[0056] (1) The magnesium-based medical material provided by the present invention significantly improves the bonding strength between the coating and the magnesium alloy substrate. By introducing a Zn metal transition layer on the surface of the magnesium alloy, the present invention can form a stable metal interface between the magnesium alloy substrate and the HA ceramic coating, thereby effectively improving the problem of insufficient adhesion of traditional ceramic coatings, reducing the phenomenon of coating peeling or detachment during use, and improving the overall stability of the coating.

[0057] (2) The magnesium-based medical material provided by the present invention effectively improves the corrosion resistance of magnesium alloy. The present invention utilizes the Zn layer to form the first protective barrier on the surface of magnesium alloy, which slows down the corrosion of the substrate by the corrosive medium. At the same time, the ZnO@HA composite layer further enhances the shielding effect of the coating, thereby forming a multi-layer protective structure and significantly reducing the corrosion rate of magnesium alloy in the corrosive environment.

[0058] (3) The magnesium-based medical material provided by this invention endows the material with good bioactivity. The HA layer of this invention can simulate the mineral composition of natural bone tissue, providing a good adhesion environment for cells, which is conducive to cell growth and bone tissue formation. Therefore, this composite coating can significantly improve the application potential of magnesium alloy materials in the biomedical field.

[0059] (4) The magnesium-based medical material provided by the present invention has a certain antibacterial ability. The present invention utilizes the Zn / ZnO structure to release Zn with antibacterial effect. 2+ The ions inhibit bacterial growth and reproduction. Therefore, this composite coating not only has good biocompatibility but also reduces the risk of bacterial infection caused by implanted materials to a certain extent.

[0060] (5) The magnesium-based medical material provided by the present invention can improve the stability of the coating structure. The present invention utilizes the flexible properties of the PCL biodegradable polymer layer to form a continuous and dense protective layer on the coating surface, thereby buffering the influence of external stress on the ceramic coating and effectively reducing the generation and propagation of cracks, significantly improving the crack resistance and structural stability of the coating.

[0061] (6) By constructing a multi-layer composite structure consisting of a metal layer, a ceramic layer and a polymer layer, the present invention enables the coating to simultaneously possess corrosion resistance, bioactivity, antibacterial properties and good structural stability, thereby giving the obtained Zn / ZnO@HA / PCL bio-coating a more complete comprehensive performance.

[0062] (7) Regulation of magnesium alloy degradation rate: The composite coating of this invention can regulate the degradation rate of magnesium alloy in body fluid environment to a certain extent, making its degradation process more stable and controllable, thereby improving the safety and reliability of the material in the biomedical field. The regulation of magnesium alloy degradation rate is essentially a synergistic regulation of interfacial electrochemical reaction, corrosion product evolution and ion transport behavior. Among them, the Zn transition layer effectively alleviates the potential difference between the magnesium matrix and the ceramic layer, inhibits microgalvanic corrosion and improves interfacial bonding stability; the ZnO@HA layer provides Ca 2+ PO4 3- and Zn 2+ Bioactive ions induce bone-like apatite deposition during corrosion, achieving dynamic repassivation and functionalization of the interface; the outer PCL layer, relying on its excellent film-forming properties and hydrophobicity, significantly reduces the electrolyte permeation rate and delays Cl- degradation. - PCL effectively inhibits initial rapid corrosion by eroding the micro- and nano-pores. Simultaneously, its capillary penetration and physical sealing effect further reduces localized corrosion sensitivity, promoting a shift from localized pitting corrosion to uniform degradation. With prolonged immersion time, PCL gradually hydrolyzes, achieving controlled exposure of the inorganic functional layer. This, combined with ion release and interfacial mineralization, constructs a phased degradation mechanism consisting of initial barrier inhibition, mid-term stable degradation, and late-stage release of bioactive substances.

[0063] (8) The preparation method of the present invention is simple and has good repeatability. The present invention uses Zn deposition, hydrothermal reaction and spin coating to prepare composite coatings. The preparation process is relatively simple, the operating conditions are easy to control, and it has good repeatability and practical application potential.

[0064] In summary, this invention achieves synergistic effects among the metal, ceramic, and polymer layers by constructing a multilayer composite bio-coating structure on the surface of a magnesium alloy substrate, consisting of a Zn transition layer, a ZnO@HA functional layer, and a PCL flexible protective layer. This structure not only significantly improves the bonding strength between the HA coating and the substrate but also effectively overcomes the inherent brittleness of bioceramic coatings. Simultaneously, it endows the material with excellent antibacterial properties and biocompatibility. Therefore, while improving the corrosion resistance of the magnesium alloy, it ensures the material's mechanical support capacity during bone tissue repair, meeting the modern medical demand for multifunctional orthopedic implant materials. Attached Figure Description

[0065] Figure 1The image shows the surface morphology of the Zn / ZnO@HA / PCL composite coating prepared in Example 1. As can be seen from the image, the Zn / ZnO@HA / PCL composite coating surface is uniform and continuous, without obvious cracks or macroscopic defects, indicating good film quality and structural integrity. During spin coating, PCL effectively penetrates and fills the micro- and nano-pores of the ZnO@HA layer, achieving pore sealing and densification, while retaining appropriate surface undulations, which is beneficial for balancing corrosion resistance and bioactivity.

[0066] Figure 2 The XRD patterns of the Zn / ZnO@HA / PCL composite coating prepared in Example 1, the Zn / ZnO@HA composite coating prepared in Comparative Example 4, the PCL coating prepared in Comparative Example 3, and the AZ31B magnesium alloy substrate are shown in the figures. As can be seen from the figures, the AZ31B substrate mainly exhibits characteristic diffraction peaks of Mg, while characteristic peaks of Zn, ZnO, and HA can be observed in the Zn / ZnO@HA coating, indicating that the composite inorganic layer has been successfully constructed and the structures of each phase are stable. After further introducing PCL, no obvious new crystal peaks were observed in the Zn / ZnO@HA / PCL and AZ31B / PCL samples; only background enhancement or broad diffuse peaks were observed, indicating that PCL exists in an amorphous state and has not changed the crystal structure of the original inorganic phase.

[0067] Figure 3 The images show the FTIR spectra of the Zn / ZnO@HA / PCL composite coating prepared in Example 1, the Zn / ZnO@HA composite coating prepared in Comparative Example 4, and the PCL coating prepared in Comparative Example 3. As can be seen from the figures, PO42- is present in both the Zn / ZnO@HA and Zn / ZnO@HA / PCL samples. 3- Characteristic absorption peaks (e.g., ~560, 600 and ~1030 cm⁻¹) -1 This indicates that HA was successfully constructed; while in the PCL-containing samples (Zn / ZnO@HA / PCL and AZ31B / PCL), ~1720 cm⁻¹ -1 The C=O stretching vibration peak and ~2940 cm⁻¹ -1 The presence of a prominent CH2- characteristic peak nearby indicates that PCL has been successfully introduced and formed an organic layer. In contrast, organic and inorganic characteristic peaks coexist in Zn / ZnO@HA / PCL, and the intensity of some peaks changes, indicating that the PCL-ZnO@HA composite structure has formed and plays a regulatory role in the surface chemical environment.

[0068] Figure 4 The figures show the Tafel curves of the Zn / ZnO@HA / PCL composite coating prepared in Example 1, the PCL coating prepared in Comparative Example 3, the Zn / ZnO@HA coating prepared in Comparative Example 4, and the AZ31B magnesium alloy in SBF solution. As can be seen from the figures, the Et of the Zn / ZnO@HA / PCL composite coating... corrClearly positive shift and i corr The significant reduction indicates that it has the best corrosion inhibition ability on the AZ31B substrate. In comparison, although Zn / ZnO@HA and AZ31B / PCL can both reduce the corrosion rate to some extent, the former is limited by its microporous structure and the latter lacks a stable interfacial transition layer, so their protective effects are not as good as those of the composite coatings. The bare AZ31B exhibits the lowest E0.05. corr and the highest i corr Corrosion is most severe there.

[0069] Figure 5 The image shows the surface morphology of the Zn / ZnO@HA / PCL composite coating prepared in Example 1 after immersion in SBF solution for 21 days. As can be seen from the image, the Zn / ZnO@HA / PCL composite coating maintained good surface integrity after immersion in SBF solution for 21 days, without obvious through cracks or large-area peeling, only uniformly distributed deposits covering the surface. Simultaneously, bone-like apatite particles were observed to form on the surface and gradually densify, indicating that the coating effectively inhibits corrosion while inducing biomineralization, demonstrating good interfacial stability and bioactivity.

[0070] Figure 6 The MgO substrate prepared by Zn / ZnO@HA / PCL in Example 1, the PCL coating prepared in Comparative Example 3, the Zn / ZnO@HA coating prepared in Comparative Example 4, and the AZ31B substrate were immersed in SBF solution for 21 days. 2+ Release line graph. As can be seen from the graph, the AZ31B sample exhibited the highest Mg content throughout the entire immersion process. 2+ The release rate indicates that it undergoes rapid and uncontrollable corrosion in SBF solution. In contrast, the AZ31B / PCL sample showed some inhibition initially, but its protective effect gradually weakened with prolonged immersion time. The Zn / ZnO@HA coating can reduce Mg content to some extent. 2+ The release rate is mainly attributed to the interfacial stabilizing effect of the Zn transition layer and the HA-induced redeposition behavior. However, due to its porous structure, it is still difficult to completely block the penetration of corrosive media. The Zn / ZnO@HA / PCL composite coating exhibits the lowest and most stable Mg content throughout the entire immersion cycle. 2+ The release rate indicates that it can effectively inhibit initial corrosion and achieve long-term stable degradation behavior. This is mainly due to the dense barrier effect of the PCL layer and its sealing effect on micro-nano pores, combined with the bioactivity and repassivation function of the ZnO@HA layer, which together achieve the staged regulation of the degradation rate of magnesium alloy.

[0071] Figure 7The images show scratch optical images of the Zn / ZnO@HA / PCL coating in Example 1, the PCL coating prepared in Comparative Example 3, and the Zn / ZnO@HA coating prepared in Comparative Example 4. As can be seen from the images, the Zn / ZnO@HA / PCL composite coating has clear boundaries and no obvious peeling, only slight material accumulation, indicating that its dominant mechanism is elastoplastic deformation, resulting in high interfacial bonding strength and structural stability. In contrast, the Zn / ZnO@HA coating exhibits significant crack propagation and localized peeling during scratching, displaying typical brittle fracture characteristics, indicating limited damage resistance. The AZ31B / PCL coating mainly shows significant plastic flow and material accumulation; although there are no obvious cracks, it is easily scratched, indicating low mechanical load-bearing capacity.

[0072] Figure 8 The histograms show the mechanical properties of the Zn / ZnO@HA / PCL coating in Example 1, the PCL coating prepared in Comparative Example 3, and the Zn / ZnO@HA coating prepared in Comparative Example 4. The figures show that the Zn / ZnO@HA / PCL composite coating exhibits the highest critical load (LC), indicating excellent interfacial bonding strength and anti-stripping ability. Simultaneously, its plasticity index (PI) is significantly increased, while the elastic recovery rate (R) and surface stress (σ) are significantly decreased, suggesting that the coating tends to undergo plastic energy dissipation rather than elastic recovery under applied load. These results indicate that the introduction of PCL effectively modulates the mechanical response behavior of the coating, transforming it from the typical brittle fracture mode of the ZnO@HA layer to a mechanism dominated by elastoplastic co-deformation, thereby significantly alleviating the inherent brittleness of the HA coating and improving the overall structural stability.

[0073] Figure 9 This is a schematic diagram of the bio-magnesium alloy coating material structure of Example 1. Detailed Implementation

[0074] The endpoints and any values ​​of the ranges disclosed herein are not limited to the precise ranges or values, and these ranges or values ​​should be understood to include values ​​close to these ranges or values. For numerical ranges, the endpoint values ​​of the various ranges, the endpoint values ​​of the various ranges and individual point values, and individual point values ​​can be combined with each other to obtain one or more new numerical ranges, which should be considered as specifically disclosed herein.

[0075] The present invention will be further described below with reference to specific embodiments, but the scope of protection of the present invention is not limited to the following specific embodiments. Obviously, the embodiments described below are only a part of the embodiments, and all other embodiments obtained by those skilled in the art without creative effort are still within the scope of protection of the present invention.

[0076] Unless otherwise specified, all raw materials, reagents, instruments and equipment used in this invention can be purchased from the market or prepared by existing methods.

[0077] Example 1

[0078] (1) Substrate pretreatment

[0079] After the AZ31B magnesium alloy sheet is cut into samples of the appropriate size, it is coarsely ground and then finely ground using CW sandpaper (600~1500#). After grinding, it needs to be finely ground with professional metallographic wet sandpaper (2000#) to achieve a smooth surface. On this basis, it is polished to a mirror finish in an anhydrous ethanol environment. Then, it is cleaned in an ultrasonic cleaner for 15 minutes, dried with a hair dryer on cold air, and vacuum-sealed for later use.

[0080] (2) Preparation of Zn transition layer

[0081] Sputtering was performed before deposition, with negative bias voltages set sequentially to -800, -600, -400, and -200V, each lasting 40 seconds, to sputter the magnesium alloy substrate and remove surface impurities. A 99.99% pure Zn target was then used as the arc source, and the deposition was performed when the vacuum level was less than 4 × 10⁻⁶ V. -3 The magnetically filtered cathode vacuum arc deposition was initiated at Pa, with the following settings: arc initiation current set to 90A; magnetic field current in the filter bend tube set to 2A; magnetic field current in the straight tube set to 3.5A; beam current intensity set to 400mA; negative bias voltage set to 100V; duty cycle set to 90%; deposition time set to 15min; and vacuum environment cooling time set to 24h.

[0082] (3) Preparation of Zn / ZnO@HA coating

[0083] Preparation of hydrothermal solution: Slowly add 0.15M KH2PO4 to 0.25M Ca-EDTA solution (the atomic ratio of Ca to P is 1:1.67) while stirring. Adjust the pH of the mixed solution to 8.8 with 2.0 M NaOH.

[0084] The sample containing the zinc transition layer was immersed in a hydrothermal reaction solution and placed in a high-pressure reactor, where it was reacted at 120 °C for 15 h in a vacuum drying oven. During the hydrothermal process, ZnO was generated in situ from the Zn layer. As the reaction proceeded, an HA coating gradually grew and deposited on the Zn and ZnO surfaces, ultimately yielding a magnesium alloy containing Zn / ZnO@HA.

[0085] (4) Preparation of Zn / ZnO@HA / PCL coating

[0086] Preparation of polycaprolactone colloid: Polycaprolactone particles (number average molecular weight 45,000) were dissolved in dichloromethane solution and stirred with a magnetic stirrer for 24 h to prepare a 7.5% (mass fraction) PCL solution.

[0087] First, the pH of a mixed solution of 20 g / L KH-550 and 90% anhydrous ethanol was adjusted to 5 using glacial acetic acid. Then, the magnesium alloy containing Zn / ZnO@HA prepared in step (3) was immersed in the mixed solution for 5 min, rinsed with ultrapure water, and dried for later use. Next, polycaprolactone colloid was coated onto the surface of the magnesium alloy containing Zn / ZnO@HA using a spin-coating method. During PCL spin-coating, the rotation speed was set to 500 rpm for 60 s, then adjusted to 1000 rpm for 30 s. Finally, the mixture was cured in situ at room temperature for 24 h to obtain a magnesium-based medical material containing a Zn / ZnO@HA / PCL coating.

[0088] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 5 µm, the thickness of the ZnO@HA layer is 45 µm, and the thickness of the PCL polymer is 6 µm.

[0089] Example 2

[0090] This embodiment is carried out using a method similar to that of Embodiment 1, except that the arc initiation current of the magnetic filter cathode vacuum arc deposition in step (2) is adjusted to 120 A.

[0091] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 6 µm, the thickness of the ZnO@HA layer is 46.5 µm, and the thickness of the PCL polymer is 6 µm.

[0092] Setting the arc initiation current too high can adversely affect coating performance. An excessively high arc initiation current leads to excessive energy input in the early stages of deposition, causing localized overheating of the substrate. This results in unstable coating growth, making it prone to defects such as porosity, microcracks, and coarse grains. At the same time, high-energy particle bombardment may also damage the already formed initial dense layer, reducing the coating's density and interfacial bonding strength, ultimately leading to a decline in its corrosion resistance and mechanical properties.

[0093] Example 3

[0094] This embodiment is carried out using a method similar to that of Embodiment 1, except that the negative bias voltage for the magnetic filter cathode vacuum arc deposition in step (2) is set to 80 V.

[0095] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 4 µm, the thickness of the ZnO@HA layer is 42 µm, and the thickness of the PCL polymer is 6 µm.

[0096] Excessively low negative bias voltage during deposition weakens the bombardment effect of ions on the substrate, reducing the surface atomic migration ability and thus decreasing the density of the coating, making it prone to forming a loose structure or porosity defects. At the same time, the "cleaning effect" and ion implantation at the interface are weakened, which is not conducive to improving the bonding strength between the coating and the substrate, and may reduce the mechanical properties and corrosion resistance of the coating.

[0097] Example 4

[0098] This embodiment is carried out using a method similar to that of Embodiment 1, except that the deposition time of the magnetic filter cathode vacuum arc deposition in step (2) is adjusted to 5 min.

[0099] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 2µm, the thickness of the ZnO@HA layer is 40µm, and the thickness of the PCL polymer is 6µm.

[0100] Insufficient deposition time leads to inadequate coating thickness, making it difficult to form a continuous and dense cover layer. This results in partial exposure of the substrate or coverage by only a weak layer, thus reducing the coating's protective shielding effect. Simultaneously, incomplete coating growth results in underdeveloped grains, poor structural density, and a predisposition to porosity or defects. Furthermore, the failure to establish an effective transition layer or diffusion bonding at the interface leads to low bonding strength between the coating and the substrate, ultimately hindering the improvement of its corrosion resistance and mechanical properties.

[0101] Example 4

[0102] This embodiment is carried out in a similar manner to Embodiment 1, except that the duty cycle of the magnetic filter cathode vacuum arc deposition in step (2) is adjusted to 100%.

[0103] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 6 µm, the thickness of the ZnO@HA layer is 47 µm, and the thickness of the PCL polymer is 6 µm.

[0104] An excessively high duty cycle leads to a prolonged plasma duration per unit time, subjecting the substrate and coating surfaces to higher average energy input. This can easily cause excessively high local temperatures, resulting in instability during coating growth. Excessive heat input may promote grain coarsening and even induce internal stress accumulation, leading to defects such as microcracks or peeling in the coating. Simultaneously, continuous and intense ion bombardment can also produce a "re-sputtering" effect on the already deposited layer, disrupting its dense structure and thus reducing the coating's bonding strength and corrosion resistance.

[0105] Example 6

[0106] This embodiment is carried out using a method similar to that of Embodiment 1, except that the pH value of the hydrothermal solution in step (3) is adjusted to 6.

[0107] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 5 µm, the thickness of the ZnO@HA layer is 35 µm, and the thickness of the PCL polymer is 6 µm.

[0108] A low pH in the solution before the hydrothermal reaction significantly affects the form of ions and the nucleation process in the reaction system. Under acidic conditions, H+ in the solution... + Higher concentrations will inhibit OH − The formation of these substances hinders the nucleation and growth of products such as ZnO and HA, leading to a reduced deposition rate. Simultaneously, an overly acidic environment may corrode or dissolve the substrate or the existing precursor layer, compromising interfacial stability. Furthermore, low pH conditions result in poor crystal growth orientation, easily leading to products with uneven size or loose structure, ultimately reducing the coating's density, bonding strength, and its biological / corrosion resistance properties.

[0109] Example 7

[0110] This embodiment is carried out using a method similar to that of Embodiment 1, except that the temperature of the hydrothermal reaction in step (3) is adjusted to 90 °C, while the time remains unchanged.

[0111] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 5 µm, the thickness of the ZnO@HA layer is 38 µm, and the thickness of the PCL polymer is 6 µm.

[0112] Excessively low hydrothermal reaction temperatures lead to insufficient system reaction kinetics, reduced ion diffusion rates and reactivity, thereby inhibiting crystal nucleation and growth, resulting in a low deposition rate. Simultaneously, lower temperatures are unfavorable for the orderly arrangement of crystal structures, easily forming particles with low crystallinity, small size, and uneven distribution, resulting in a loose and insufficiently dense overall coating structure. Furthermore, insufficient interfacial reactions under low-temperature conditions make it difficult to form a stable interfacial bonding layer, thus reducing the bonding strength between the coating and the substrate, as well as its corrosion resistance and biocompatibility.

[0113] Example 8

[0114] This embodiment is carried out using a method similar to that of Embodiment 1, except that the hydrothermal reaction time in step (3) is adjusted to 8 h, while the temperature remains unchanged.

[0115] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 5 µm, the thickness of the ZnO@HA layer is 30 µm, and the thickness of the PCL polymer is 6 µm.

[0116] Insufficient hydrothermal reaction time leads to incomplete reaction, with crystal nucleation and growth processes not yet completed, resulting in insufficient product deposition and difficulty in forming a continuous and dense coating structure. Furthermore, incomplete crystal development results in low crystallinity, small and unevenly distributed particles, easily introducing pores or structural defects. In addition, the reaction and bonding processes at the interface are not fully established, leading to weak bonding strength between the coating and the substrate, thus affecting its corrosion resistance and bioactivity.

[0117] Example 9

[0118] Steps (1) to (3) are the same as in Example 1;

[0119] Step (4): Polycaprolactone colloid is the same as in Example 1;

[0120] The magnesium alloy containing Zn / ZnO@HA prepared in step (3) was directly coated with polycaprolactone colloid onto the surface of the magnesium alloy coating containing Zn / ZnO@HA by spin coating (without pretreatment using a mixed solution containing silane coupling agent and ethanol). The parameters of PCL spin coating were the same as in Example 1.

[0121] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 5 µm, the thickness of the ZnO@HA layer is 45 µm, and the thickness of the PCL polymer is 5 µm.

[0122] Failure to treat the intermediate coating with a silane coupling agent before spin coating PCL results in a lack of effective chemical bonding and bridging at the interface. Due to the weak polarity of the PCL molecular chain, its interaction with inorganic phases (such as Zn / ZnO@HA) relies primarily on physical adsorption, leading to poor interfacial compatibility and a tendency for interfacial debonding. Furthermore, the poor surface wettability without coupling agent modification hinders the uniform spreading and penetration of the PCL solution, easily forming interfacial voids or micro-defects. This reduces the coating's bonding strength and overall density, ultimately affecting its mechanical properties and corrosion resistance.

[0123] Example 10

[0124] This embodiment is carried out using a method similar to that of Example 1, except that the concentration of polycaprolactone colloid in step (4) is adjusted to 15 wt%.

[0125] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 5 µm, the thickness of the ZnO@HA layer is 45 µm, and the thickness of the PCL polymer is 3 µm.

[0126] During spin coating, excessively high PCL solution concentration significantly increases solution viscosity, reducing its fluidity and spreading ability. This makes it difficult for the coating to distribute evenly on the substrate surface, easily leading to uneven thickness or localized buildup. Simultaneously, higher concentrations hinder rapid solvent evaporation during spin coating, potentially causing solvent retention within the coating and resulting in porosity or defects. Furthermore, an excessively thick PCL layer weakens its interfacial bonding with underlying layers (such as the ZnO@HA layer) and may reduce the overall density and stability of the structure, thereby affecting the coating's mechanical properties and corrosion resistance.

[0127] Example 11

[0128] This embodiment is carried out in a similar manner to that of embodiment 1. The difference is that the parameters of the spin coating method in step (4) are adjusted. Specifically, the rotation speed is set to 500 rpm and held for 60 s. Then, the rotation speed is adjusted to 2000 rpm and held for 30 s.

[0129] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 5 µm, the thickness of the ZnO@HA layer is 45 µm, and the thickness of the PCL polymer is 4 µm.

[0130] During spin coating, excessively high rotation speeds can cause the solution to be rapidly ejected from the substrate surface before it has fully spread and penetrated the pores, making it difficult to form a continuous and dense polymer film. Simultaneously, high shear and rapid solvent evaporation can lead to flow instability and internal stress accumulation, further resulting in defects such as pinholes, streaks, and localized cracking in the coating. For Zn / ZnO@HA micro / nano structures, excessively high rotation speeds can also significantly weaken the capillary penetration and pore-sealing effect of PCL on the micro / nano pores, reducing interfacial adhesion and structural integrity, ultimately decreasing the protective performance of the coating and making it difficult to effectively inhibit the initial corrosion behavior of magnesium alloys.

[0131] Example 12

[0132] This embodiment is carried out in a similar manner to that of embodiment 1. The difference is that the parameters of the spin coating method in step (4) are adjusted. Specifically, the rotation speed is set to 500 rpm and held for 2 minutes. Then, the rotation speed is adjusted to 1000 rpm and held for 3 minutes.

[0133] In this embodiment, the thickness of the Zn transition layer of the magnesium-based medical material with Zn / ZnO@HA / PCL coating is 5 µm, the thickness of the ZnO@HA layer is 45 µm, and the thickness of the PCL polymer is 2 µm.

[0134] Prolonged spin coating accelerates solvent evaporation, potentially preventing sufficient polymer chain rearrangement during film formation and affecting the coating's density and uniformity. An excessively thin coating also weakens its protective effect on the underlying structure, negatively impacting overall corrosion resistance and mechanical stability.

[0135] Comparative Example 1

[0136] This comparative example was performed using a method similar to that of Example 1, except that steps (3) and (4) were omitted, and only a magnesium-based medical material containing a Zn transition layer was obtained. The thickness of the Zn transition layer was 5 µm.

[0137] Electrochemical test results show that the E of the exposed AZ31B magnesium alloy corr -1.58 V, i corr It is 2.41 × 10 -4 A / cm 2 The sample treated with Zn plating showed a positive shift in corrosion potential to -1.40 V, and a significant reduction in corrosion current density to 1.60 × 10⁻⁶ V. -5 A / cm 2Compared to the substrate material, the corrosion potential of the Zn-plated sample shifted significantly to the positive side, indicating a reduced thermodynamic corrosion tendency. Simultaneously, the corrosion current density decreased by approximately one order of magnitude, suggesting that the corrosion reaction kinetics were effectively suppressed. These results indicate that Zn provides electrochemical protection to the substrate to a certain extent, thereby significantly improving the corrosion resistance of the AZ31B magnesium alloy. However, a single Zn layer is insufficient to halt the corrosion process, which also means that its corrosion resistance cannot match that of the composite coating in Example 1.

[0138] Comparative Example 2

[0139] Step (1) is the same as in Example 1;

[0140] Step (2): Preparation of hydrothermal solution: Slowly add 0.15M KH2PO4 to 0.25M Ca-EDTA solution (the atomic ratio of Ca to P is 1:1.67) while stirring. Adjust the pH of the mixed solution to 8.8 with 2.0M NaOH.

[0141] The pretreated magnesium alloy sample was immersed in a hydrothermal reaction solution and placed in a high-pressure reactor, where it was reacted at 120°C for 15 hours in a vacuum drying oven. During the hydrothermal process, an HA coating gradually grew and deposited on the surface of the magnesium alloy, ultimately yielding a magnesium-based medical material containing only the HA layer. The thickness of the HA layer was 42 µm.

[0142] Directly depositing HA coatings on the surface of AZ31B magnesium alloys typically yields unsatisfactory interfacial bonding performance and long-term stability. Electrochemical testing results indicate that HA coatings deposited on a Zn-plated interlayer exhibit superior EH performance. corr The corrosion current density was -0.55 V, significantly higher than the -0.82 V of HA deposited directly on the AZ31B magnesium alloy surface, indicating a significant reduction in its corrosion tendency. Simultaneously, its corrosion current density was reduced by approximately three orders of magnitude compared to the control sample, indicating that the corrosion reaction rate was effectively suppressed and the corrosion resistance was significantly improved. On one hand, the magnesium alloy surface has high chemical activity, making it prone to rapid corrosion or the formation of a loose oxide / hydrogen layer during deposition, affecting the uniform nucleation and adhesion of HA. On the other hand, HA, as a ceramic phase, differs significantly from the metal substrate in its coefficient of thermal expansion and mechanical properties, easily leading to interfacial stress concentration, thereby causing coating cracking or peeling. Furthermore, directly deposited HA coatings typically have a loose structure and high porosity, making it difficult to form an effective dense barrier. Corrosive media can easily penetrate into the substrate through defects, reducing its corrosion resistance. Therefore, the coating prepared according to Comparative Example 2 cannot match the performance of the composite coating in Example 1.

[0143] Comparative Example 3

[0144] Step (1) is the same as in Example 1;

[0145] Step (2): Preparation of polycaprolactone colloid: Dissolve polycaprolactone particles (number average molecular weight of 45,000) in dichloromethane solution and stir with a magnetic stirrer for 24 h to prepare a 7.5% (mass fraction) PCL solution.

[0146] First, the pH of a mixed solution of 20 g / L KH-550 and 90% anhydrous ethanol was adjusted to 5 using glacial acetic acid. Then, the pretreated magnesium alloy plate sample was immersed in the mixed solution for 5 min, rinsed with ultrapure water, and dried. Next, polycaprolactone (PCL) colloid was spin-coated onto the magnesium alloy surface. During PCL spin-coating, the rotation speed was set to 500 rpm for 60 s, then adjusted to 1000 rpm for 30 s. Finally, the mixture was cured in situ at room temperature for 24 h to obtain a magnesium-based medical material containing only a PCL coating. The thickness of the PCL polymer was 5 µm.

[0147] Depend on Figure 6 It can be seen that during the degradation process in simulated body fluid (SBF) at 37 ℃, the PCL sample directly spin-coated on the surface of AZ31B magnesium alloy exhibited lower Mg content. 2+ The release amount gradually increased over time, and the release rate accelerated, indicating that the single PCL coating was insufficient to effectively inhibit substrate corrosion and had limited protective effect. In contrast, after further spin-coating PCL onto the Zn / ZnO@HA composite coating surface, Mg... 2+ The release process was noticeably gradual, with a low overall release amount and stable variation, indicating that the composite coating system can effectively delay the penetration of corrosive media and achieve controllable regulation of the substrate degradation behavior. In terms of mechanical properties, the adhesion of the directly spin-coated PCL sample was only 5.58 N, significantly lower than the 13.43 N of the Zn / ZnO@HA / PCL composite coating (see...). Figures 7-9 The results indicate that the multilayer structure design significantly enhances the interfacial bonding strength between the coating and the substrate. Furthermore, the plasticity index (PI), elastic recovery rate (R), and surface stress (σ) of the Zn / ZnO@HA / PCL composite coating are 0.77, 0.27, and 0.46, respectively. The high plasticity index, low elastic recovery rate, and low surface stress indicate that this composite coating possesses excellent energy dissipation and deformation coordination capabilities, thus exhibiting strong toughness and effectively overcoming the inherent brittleness of traditional HA coatings.

[0148] Comparative Example 4

[0149] This comparative example was performed using a method similar to that of Example 1, except that step (4) of Example 1 was omitted, and only a magnesium-based medical material containing Zn / ZnO@HA was obtained. The thickness of the Zn transition layer was 5 µm, and the thickness of the ZnO@HA coating was 45 µm.

[0150] Compared to constructing only a Zn / ZnO@HA inorganic coating, introducing a PCL outer layer significantly improves the overall protective performance and interfacial stability of the coating. Without PCL modification, the ZnO@HA layer, due to its inherent porous structure, is unable to effectively block electrolyte penetration, leading to rapid arrival of corrosive media in the substrate and inducing localized pitting corrosion, thus accelerating the initial degradation process of the magnesium alloy. Although the ZnO@HA layer can achieve some degree of repassivation by inducing bone-like apatite deposition, this process usually lags behind corrosion occurrence and is insufficient to suppress early rapid corrosion behavior. In contrast, the PCL layer, by constructing a dense barrier and sealing micro- and nano-pores, effectively delays the transport of corrosive media while enabling controlled ion release, thereby transforming the corrosion mode from localized pitting corrosion to a relatively uniform degradation process, significantly improving the long-term stability and synergistic biological functions of the system.

[0151] Comparative Example 5

[0152] Step (1) is the same as in Example 1;

[0153] Step (2): HA coating preparation: prepare hydrothermal solution by slowly adding 0.15M KH2PO4 to 0.25M Ca-EDTA solution (the atomic ratio of Ca to P is 1:1.67) while stirring. Adjust the mixed solution to 8.8 with 2.0M NaOH.

[0154] The pretreated sample obtained in step (1) was immersed in a hydrothermal reaction solution and placed in a high-pressure reactor, where it was reacted at 120 °C for 15 h in a vacuum drying oven. During the hydrothermal process, an HA coating was gradually grown and deposited, ultimately yielding a magnesium alloy containing an HA layer.

[0155] Step (3): Preparation of HA / PCL coating

[0156] Preparation of polycaprolactone colloid: Polycaprolactone particles (number average molecular weight 45,000) were dissolved in dichloromethane solution and stirred with a magnetic stirrer for 24 h to prepare a 7.5% (mass fraction) PCL solution.

[0157] First, the pH of a mixed solution of 20 g / L KH-550 and 90% anhydrous ethanol was adjusted to 5 using glacial acetic acid. Then, the magnesium alloy containing the HA layer prepared in step (2) was immersed in the mixed solution for 5 min, rinsed with ultrapure water, and dried for later use. Then, polycaprolactone colloid was coated onto the surface of the magnesium alloy coating containing the HA layer using spin coating. During PCL spin coating, the rotation speed was set to 500 rpm and continued for 60 s. After that, the rotation speed was adjusted to 1000 rpm and continued for 30 s. Finally, the magnesium-based medical material containing the HA / PCL coating was obtained by in-situ curing at room temperature for 24 h.

[0158] In this comparative example, the thickness of the HA coating is 42 µm, and the thickness of the PCL polymer is 5 µm.

[0159] While directly hydrothermally generating HA on the AZ31B surface and then coating it with PCL can construct an organic / inorganic composite coating system, which can improve surface bioactivity and delay the corrosion process to some extent, the lack of an effective interfacial transition layer between the magnesium substrate and HA results in weak interfacial adhesion and susceptibility to failure during corrosion. Furthermore, the pre-corrosion of the magnesium substrate during hydrothermal processing and the porous structure of the HA layer itself further exacerbate the risk of corrosive media penetration. Although the PCL layer can improve surface density to some extent through pore sealing, it is difficult to fundamentally suppress interfacial corrosion and coating peeling.

[0160] Test case

[0161] The performance of each coating sample prepared in the above examples was tested, specifically including corrosion potential (E). corr ), corrosion current density (i corr ), surface stress (σ), bonding force (LC), elastic recovery rate (R), and plasticity index (PI). Corrosion potential (E) is also considered. corr ), corrosion current density (i corr The testing standard for the surface stress (σ), bonding strength (LC), elastic recovery rate (R), and plasticity index (PI) is ASTM G5-14. The surface stress (σ), bonding strength (LC), elastic recovery rate (R), and plasticity index (PI) are calculated using nano-scratching, and the testing standard for nano-scratching is ASTM C1624. Specific test results are shown in Table 1.

[0162]

[0163] As can be seen from the experimental data in Table 1, the Zn-ZnO@HA / PCL composite coating, through the synergistic effect of "in-situ synthesis + inorganic functional layer + organic flexible layer", significantly improves corrosion resistance while simultaneously optimizing mechanical properties, structural stability, antibacterial properties, and biocompatibility.

[0164] The preferred embodiments of the present invention have been described in detail above; however, the present invention is not limited thereto. Within the scope of the inventive concept, various simple modifications can be made to the technical solutions of the present invention, including combinations of various technical features in any other suitable manner. These simple modifications and combinations should also be considered as the content disclosed in the present invention and are all within the protection scope of the present invention.

Claims

1. A magnesium-based medical material containing a Zn / ZnO@HA / PCL bio-coating, characterized in that: The magnesium alloy-based medical material includes a magnesium alloy matrix and a Zn transition layer, a ZnO@HA layer and a PCL polymer layer sequentially stacked on the magnesium alloy matrix; ZnO@HA is a hydroxyapatite-zinc oxide composite material.

2. The magnesium-based medical material containing a Zn / ZnO@HA / PCL bio-coating according to claim 1, characterized in that: The magnesium alloy substrate is selected from at least one of WE43 magnesium alloy, AZ31B magnesium alloy, ZK60 magnesium alloy, and pure magnesium; Preferably, the magnesium alloy substrate is AZ31B magnesium alloy.

3. A magnesium-based medical material containing a Zn / ZnO@HA / PCL bio-coating according to claim 1 or 2, characterized in that: The thickness of the Zn transition layer is 1~10 µm; And / or, the thickness of the ZnO@HA layer is 20~50 µm; And / or, the thickness of the PCL polymer is 5~10 µm.

4. A method for preparing a magnesium-based medical material containing a Zn / ZnO@HA / PCL bio-coating as described in claims 1-3, characterized in that: A Zn transition layer is formed sequentially on a magnesium alloy substrate, followed by a ZnO@HA layer and a PCL polymer layer.

5. The method for preparing a magnesium-based medical material containing a ZnO@HA / PCL bio-coating according to claim 4, characterized in that: Includes the following steps: (1) A Zn transition layer is formed on a magnesium alloy substrate by physical vapor deposition to obtain a magnesium alloy containing a Zn transition layer; (2) A magnesium alloy containing a Zn transition layer is placed in a hydrothermal solution and subjected to hydrothermal reaction to deposit a magnesium alloy containing Zn / ZnO@HA. (3) Polycaprolactone colloid is coated onto the surface of a magnesium alloy coating containing Zn / ZnO@HA by spin coating and cured.

6. The method for preparing magnesium-based medical materials containing a ZnO@HA / PCL bio-coating according to claim 4 or 5, characterized in that: Before the physical vapor deposition method, the magnesium alloy substrate is sputtered. Preferably, the negative bias voltage of the sputtering is adjusted sequentially to -800 ± 50 V, -600 ± 50 V, -400 ± 50 V and -200 ± 50 V, and the deposition time at each negative bias voltage point is independently 20~40 s.

7. The method for preparing magnesium-based medical materials containing a ZnO@HA / PCL bio-coating according to claim 4 or 5, characterized in that: The physical vapor deposition method is selected from one of the following: magnetically filtered cathode vacuum arc deposition, multi-arc ion plating, and magnetron sputtering. Preferably, the physical vapor deposition method is magnetically filtered cathode vacuum arc deposition. Preferably, the process of the magnetically filtered cathode vacuum arc deposition method is as follows: controlling the vacuum degree of the vacuum chamber of the magnetically filtered cathode vacuum arc deposition equipment to be less than or equal to 5 × 10⁻⁶. -3 After Pa, using a Zn target as the arc source, during the vapor deposition process, the arc initiation current was controlled at 60~120 A, the bending magnetic field current at 1.5~3.0 A, the beam current intensity at 350~400 mA, the negative bias voltage at 90~100 V, the duty cycle at 60~90%, the deposition time at 10~30 min, and the vacuum environment cooling time at >1 h.

8. The method for preparing magnesium-based medical materials containing a ZnO@HA / PCL bio-coating according to claim 4 or 5, characterized in that: The hydrothermal solution contains soluble salts containing calcium and soluble salts containing phosphate, and the atomic ratio of Ca to P is 1:1.66~1.68; And / or, the pH of the hydrothermal solution is 8.5~11, the temperature of the hydrothermal reaction is 100~150℃, and the time is 12~16 h.

9. The method for preparing magnesium-based medical materials containing a ZnO@HA / PCL bio-coating according to claim 4 or 5, characterized in that: The polycaprolactone in the polycaprolactone colloid has a number-average molecular weight of 40,000 to 60,000; the solvent in the polycaprolactone colloid is dichloromethane. And / or, the concentration of the polycaprolactone colloid is 5~10wt%, the spin coating rate is 500~1000 rpm, and the spin coating time is 30~100 s.

10. The application of a magnesium-based medical material containing a ZnO@HA / PCL bio-coating as described in any one of claims 1 to 3, or a magnesium-based medical material containing a ZnO@HA / PCL bio-coating prepared by any one of claims 4 to 9, as an orthopedic repair implant.

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