Medical degradable zinc-based material surface structure and preparation method and application thereof

Abstract

The invention discloses a medical degradable zinc-based material surface structure which comprises a zinc-based matrix, a calcium phosphate salt layer is loaded on the zinc-based matrix, and the surface of the calcium phosphate salt layer is subjected to laser remelting to form a remelting area. The surface structure of the zinc-based material is stable, the comprehensive mechanical property, especially the wear resistance and the degradation property of the zinc-based material are improved, and the zinc-based surface material can be quickly degraded and absorbed by human tissues and is beneficial to cell growth. The invention further discloses a preparation method of the medical degradable zinc-based material surface structure. The method is easy to operate, simple in process, short in practice period and capable of being put into application rapidly and massively. The surface structure of the zinc-based material can be quickly degraded and absorbed by human tissues, is combined with osteoblasts, is beneficial to cell growth and accelerates bone healing, and can be applied to preparation of bone healing materials.

Description

Technical Field

[0001] The invention relates to the technical field of metal material surface treatment, and in particular to a medical degradable zinc-based alloy surface structure and a preparation method and application thereof. Background Art

[0002] Degradable medical metal materials usually refer to a type of functional materials used in the diagnosis and repair of tissues or organs and other disease treatment fields, which have no side effects on tissues, organs and blood. With the rapid development of medical tissue engineering and materials science, the "degradability" of medical metal materials has received increasing attention in recent years.

[0003] Degradable medical metal materials represented by magnesium, zinc, and iron have attracted more and more attention from scholars in the biomedical and material fields due to their complete degradability in the body. Among them, degradable medical zinc and zinc alloy materials are expected to become ideal candidate materials for the new generation of degradable bone implants through further performance optimization due to their moderate degradation rate and mechanical properties. However, the current research on the application of zinc alloys in orthopedic materials is still in its infancy, and there is still a lack of sufficient research information and clinical data support for the mechanical properties, degradation properties, and mechanism of its osteoinductive properties. How to improve mechanical properties, control degradation rate, and improve surface osteogenic properties are common problems for zinc and zinc alloys. Ideal orthopedic implant materials need to have high strength, toughness, suitable degradation rate, and good osteogenic properties. Especially in the early stage of implantation, degradable zinc alloys need to have sufficient mechanical properties to maintain device stability and better osteogenic properties to promote bone tissue healing.

[0004] The mechanical properties of zinc alloys have been significantly improved by adding alloying elements and changing processing technology. However, a large number of studies have shown that zinc alloys have obvious degradation and wear behaviors in the physiological environment. In this process, excessive zinc ions are released, which have certain cytotoxicity and cannot meet the clinical needs of direct contact with tissues. Surface modification is a common means of changing the surface properties of implant materials. For zinc alloys, it is necessary to improve the surface mechanical properties and biocompatibility of the material, especially the osteogenic properties, through surface modification.

[0005] Publication No. CN113289071A discloses an implantable zinc-based alloy surface structure, including a substrate, and a porous zinc oxide layer formed on the substrate surface, and a calcium phosphate layer is loaded on the surface of the porous zinc oxide layer. However, the metal surface after micro-arc oxidation treatment usually forms a porous and rough ceramic layer. This structure may make it difficult for the calcium and phosphorus components in the liquid phase to deposit evenly, and it will be difficult to form a calcium phosphate layer in micropores or recesses, resulting in incomplete coverage; and the porous zinc oxide layer may be chemically incompatible with the formation reaction of the calcium phosphate layer. The chemical composition of the oxide layer may be damaged during the formation of the calcium phosphate layer, affecting the integrity of the calcium phosphate layer.

[0006] Publication No. CN117966267A discloses a zinc-based alloy surface structure containing a calcium-phosphorus biological coating, including a substrate and a calcium-phosphorus salt layer formed on the substrate surface. However, the zinc-based alloy surface structure only forms a calcium-phosphorus salt layer directly on the substrate surface by a hydrothermal method, but the bonding force between the calcium-phosphorus salt layer and the substrate is poor, and the calcium-phosphorus salt layer cannot be completely spread on the substrate, and the calcium-phosphorus layer far away from the zinc substrate end is not chemically bonded to the metal substrate. If the bonding between the calcium-phosphorus salt layer and the metal substrate is not strong enough, it will affect the stability and durability of the implant, and thus affect the bone healing process. Summary of the invention

[0007] In order to solve the above technical problems, the present invention provides a medical degradable zinc-based material surface structure, in which a calcium-phosphate layer is evenly loaded on the surface of the zinc-based material through laser remelting. The surface structure of the zinc-based material is stable, and the comprehensive mechanical properties (especially wear resistance) and degradation performance of the zinc-based material are improved. In addition, the zinc-based surface material can be degraded and absorbed by human tissues more quickly, which is beneficial to cell growth.

[0008] A medical degradable zinc-based material surface structure comprises a zinc-based matrix, a calcium phosphate layer is loaded on the zinc-based matrix, and the surface of the calcium phosphate layer is laser remelted to form a remelted area.

[0009] In the present invention, a relatively rough remelting zone is formed by laser remelting the surface of the calcium phosphate layer, and the remelting zone increases the contact area and helps to uniformly corrode. After considering the comprehensive mechanical properties and biocompatibility, the present invention loads the calcium phosphate layer on the substrate, the calcium phosphate layer is in full contact with the substrate, has a certain binding force, is wrapped by the substrate during the laser remelting process, and forms a calcium phosphate-matrix mixed remelting zone, so that the calcium phosphate layer is in full contact with the substrate and fused and evenly loaded on the substrate surface. At the same time, due to the presence of calcium phosphate, the biocompatibility of the zinc-based surface material can be improved, and it can be degraded and absorbed by human tissues more quickly.

[0010] Preferably, the zinc-based matrix is ​​zinc or a zinc alloy, and the zinc alloy is an alloy formed by zinc and at least one metal selected from manganese, iron, lithium, magnesium, copper and calcium; in the zinc alloy, 50wt%≤Zn<100wt%.

[0011] Manganese is one of the essential trace elements for the human body. It is a trace element necessary for the human immune system and various enzymes. Manganese can also promote the growth and development of human bones. Zinc-manganese alloy has good comprehensive mechanical properties. When the manganese content is below 1wt%, the zinc-manganese alloy exhibits excellent ductility at room temperature. In a specific embodiment of the present invention, the surface structure of the zinc-based material made of zinc-lithium alloy has good wear resistance and good biocompatibility. Further preferably, the zinc alloy is a zinc-manganese alloy or a zinc-lithium alloy, and the content of manganese or lithium in the zinc alloy is 0.30-0.45wt%.

[0012] Preferably, in the surface structure of the zinc-based material, 0<the thickness of the calcium phosphate salt layer is ≤20 μm, and 0<the thickness of the remelting zone is ≤300 μm.

[0013] In the present invention, the calcium phosphate layer is in full contact with the substrate. During the laser remelting process, the calcium phosphate is wrapped by the substrate to form a remelting zone. When the thickness of the remelting zone is greater than the thickness of the loaded calcium phosphate layer, the calcium phosphate in the remelting zone is in full contact with the substrate, fused, and evenly loaded on the surface of the substrate.

[0014] Further preferably, in the surface structure of the zinc-based material, the thickness of the calcium phosphate salt layer is 5 μm ≤ 10 μm, and the thickness of the remelting zone is 100 μm ≤ 120 μm.

[0015] The present invention also provides a method for preparing the surface structure of the medical degradable zinc-based material. The method has simple operation, streamlined process, short practice cycle, and can be quickly put into use in large quantities.

[0016] A method for preparing a medical degradable zinc-based material surface structure comprises the following steps:

[0017] (1) removing the oxide layer from the surface of the substrate to obtain a pretreated substrate;

[0018] (2) A calcium phosphate layer is loaded on the pretreated substrate obtained in step (1), and the surface of the calcium phosphate layer is laser remelted to obtain a medical degradable zinc-based material surface structure.

[0019] Preferably, in step (1), the oxidation layer removal treatment is to grind the substrate surface with sandpaper.

[0020] Use sandpaper to polish the substrate surface to remove surface impurities and oxide layers, making the substrate surface smooth and clean to prevent uneven heating and impurities in subsequent steps.

[0021] Preferably, in step (2), the method for loading the calcium phosphate layer is to place the pretreated substrate in a calcium phosphate precursor solution for a hydrothermal reaction.

[0022] Further preferably, the calcium-phosphorus precursor solution consists of calcium nitrate and sodium dihydrogen phosphate.

[0023] A calcium phosphate layer is prepared on the surface of the substrate by a hydrothermal method. The hydrogen phosphate ions in the calcium phosphate precursor solution combine with calcium ions to generate stable and insoluble calcium hydrogen phosphate, which is deposited and attached to the surface of the zinc-based substrate. The obtained calcium phosphate layer has a high binding force and contains elements such as Ca, P, and O. It dissolves after contacting body fluids, thereby improving biocompatibility and being degradable along with human metabolism.

[0024] More preferably, in the calcium-phosphorus precursor solution, the molar ratio of calcium nitrate to sodium dihydrogen phosphate is 1 to 2:1.

[0025] In the calcium phosphate layer, when the molar ratio of calcium nitrate to sodium dihydrogen phosphate is within the above range, the calcium phosphate layer has a higher specific surface area and a higher affinity for cells. The calcium phosphate layer dissolves after contacting with body fluids, and the dissolved calcium and phosphorus can be degraded along with human metabolism, thereby improving the biocompatibility of the zinc-based surface material.

[0026] Preferably, the temperature of the hydrothermal reaction is 100-160° C., and the reaction time is 1-8 hours.

[0027] In the present invention, the above-mentioned hydrothermal reaction conditions can improve the chemical bonding between the calcium phosphate salt layer and the zinc-based matrix, enhance the binding force of the calcium phosphate salt layer, and make the calcium phosphate salt layer evenly distributed on the surface of the zinc-based matrix, thereby ensuring that the entire surface has the same biological activity, corrosion protection ability and good biocompatibility.

[0028] Preferably, in step (2), the pulse frequency of the laser remelting is 5 to 10 Hz; and the pulse duration is 1 to 2 ms.

[0029] By adjusting the pulse frequency and pulse duration of the laser, the thickness of the remelting zone can be adjusted to obtain a remelting zone with better surface morphology and better surface mechanical strength. When the pulse frequency and pulse duration are within the above range, the matrix and calcium phosphate salt are fully combined to form a remelting zone mixed with calcium phosphate salt and matrix, and the influence of other elements doped in the air during the laser remelting process can be avoided. The high energy during the laser remelting process causes part of the Zn on the surface of the matrix to volatilize, reducing the Zn content, while reducing the element segregation in the surface structure of the zinc-based material and improving its biocompatibility.

[0030] The present invention also provides application of the above-mentioned medical degradable zinc-based material surface structure in preparing bone healing materials.

[0031] In the surface structure of the zinc-based material of the present invention, the thickness of the calcium-phosphate salt layer is 5 μm ≤ ≤ 10 μm, and the thickness of the remelting zone is 100 μm ≤ ≤ 120 μm, so that the surface structure of the zinc-based material dissolves after contacting with body fluids. The dissolved calcium and phosphorus can be degraded along with human metabolism and combined with osteoblasts, which is beneficial to cell growth and accelerates bone healing. At the same time, it can prevent the zinc-based matrix from being corroded by body fluids to release zinc ions to produce cytotoxicity.

[0032] Compared with the prior art, the present invention has the following beneficial effects:

[0033] (1) The medical degradable zinc-based material surface structure provided by the present invention is a zinc-based material surface having a calcium phosphate layer loaded thereon, and then the surface of the calcium phosphate layer is laser remelted to form a calcium phosphate-matrix mixed remelting zone. The zinc-based material surface structure is stable, and the comprehensive mechanical properties (especially wear resistance) and degradation properties of the zinc-based material are improved. In addition, the zinc-based surface material can be degraded and absorbed by human tissues more quickly, which is beneficial to cell growth.

[0034] (2) The preparation method of the surface structure of the zinc-based material provided by the present invention reduces the difficulty of combining the calcium phosphate layer with the zinc alloy matrix, replaces the traditional process, greatly shortens the practice cycle, simplifies the process, and can mass-produce the required products.

[0035] (3) The surface structure of the zinc-based material provided by the present invention can be quickly degraded and absorbed by human tissues, combined with osteoblasts, which is beneficial to cell growth, accelerates bone healing, and can be used in the preparation of bone healing materials. BRIEF DESCRIPTION OF THE DRAWINGS

[0036] Figure 1 These are surface and cross-sectional morphology images of the surface structure of the zinc-based material obtained in Example 1 and Comparative Example 1, wherein (a) and (b) are surface and cross-sectional morphology images of the surface structure of the zinc-based material obtained in Comparative Example 1, respectively; (c) and (d) are surface and cross-sectional morphology images of the surface of the zinc-based material obtained in Example 1, respectively.

[0037] Figure 2 The cross-sectional metallographic images and cross-sectional morphology images of the surface structure of the zinc-based material obtained in Example 2 and Comparative Example 2, wherein (a) and (c) are the cross-sectional metallographic images and cross-sectional morphology images of the surface structure of the zinc-based material obtained in Comparative Example 1, respectively; (b) and (d) are the cross-sectional metallographic images and cross-sectional morphology images of the surface of the zinc-based material obtained in Example 1, respectively.

[0038] Figure 3 The XRD patterns of the surface structures of the zinc-based materials obtained from the Zn-0.4Mn alloy, Example 1 and Comparative Example 1, wherein (a) is the original XRD pattern and (b) is a local enlarged view of 29° to 33° in (a).

[0039] Figure 4 The XRD patterns of the surface structures of the zinc-based materials obtained from the Zn-0.35Li alloy, Example 2 and Comparative Example 2, wherein (a) is the original XRD pattern and (b) is a local enlarged view of 28° to 32° in (a).

[0040] Figure 5 Schematic diagram of nanoindentation in surface strength testing.

[0041] Figure 6 These are surface strength test diagrams of the zinc-based material surface structures obtained in Example 1 and Comparative Example 1 under the conditions of a load of 5 μN and a duration of 10 seconds, wherein (a) and (b) are respectively the load-depth curve and the Brinell hardness change diagram of the zinc-based material surface structure obtained in Comparative Example 1, and (c) and (d) are respectively the load-depth curve and the Brinell hardness change diagram of the zinc-based material surface structure obtained in Example 1.

[0042] Figure 7 The cross-sectional microhardness distribution diagrams of the surface structures of the zinc-based materials prepared in Example 1 and Comparative Example 1, wherein (a) and (b) are the cross-sectional microhardness distribution diagrams of Comparative Example 1 and Example 1, respectively.

[0043] Figure 8 These are surface strength test diagrams of the zinc-based material surface structures obtained in Example 2 and Comparative Example 2 under the conditions of a load of 5 μN and a duration of 10 seconds, wherein (a) and (b) are respectively the load-depth curve and the Brinell hardness change diagram of the zinc-based material surface structure obtained in Comparative Example 2, and (c) and (d) are respectively the load-depth curve and the Brinell hardness change diagram of the zinc-based material surface structure obtained in Example 2.

[0044] Fig. 9 The cross-sectional microhardness distribution diagrams of the surface structures of the zinc-based materials prepared in Example 2 and Comparative Example 2, wherein (a) and (b) are the cross-sectional microhardness distribution diagrams of Comparative Example 2 and Example 2, respectively.

[0045] Fig.10 The figures are the wear resistance test diagrams of the surface structures of zinc-based materials prepared by Zn-0.4Mn alloy, Example 1 and Comparative Example 1, wherein (a) is the friction coefficient change curve, and (b) to (d) are the wear depth curves of Zn-0.4Mn alloy, Comparative Example 1 and Example 1, respectively.

[0046] Fig.11The figures are the wear resistance test diagrams of the surface structures of zinc-based materials prepared from Zn-0.35Li alloy, Example 2 and Comparative Example 2, wherein (a) is the friction coefficient change curve, and (b) to (d) are the wear depth curves of Zn-0.35Li alloy, Comparative Example 2 and Example 2, respectively.

[0047] Fig.12 The corrosion resistance test diagrams of the surface structures of zinc-based materials prepared by Zn-0.4Mn alloy, Example 1 and Comparative Example 1, wherein (a) to (c) are Nyquist curves, Bode diagrams and polarization curves, respectively.

[0048] Fig.13 These are corrosion resistance test diagrams of the surface structures of zinc-based materials prepared from Zn-0.35Li alloy, Example 2 and Comparative Example 2, wherein (a) to (c) are Nyquist curves, Bode diagrams and polarization curves, respectively.

[0049] Fig.14 It is a biocompatibility test diagram of the surface structure of zinc-based materials prepared by Zn-0.4Mn alloy, Example 1 and Comparative Example 1, wherein a to d are the test results of 25% Zn-0.4Mn alloy leaching solution group, 25% comparative example 1 leaching solution group, 25% example 1 leaching solution group and normal control group, respectively.

[0050] Fig.15 It is a biocompatibility test diagram of the surface structure of zinc-based materials prepared by Zn-0.35Li alloy, Example 2 and Comparative Example 2, wherein a to d are the test results of 25% Zn-0.35Li alloy leaching solution group, 25% Comparative Example 2 leaching solution group, 25% Example 2 leaching solution group and normal control group, respectively. DETAILED DESCRIPTION

[0051] The present invention will be further described in detail below in conjunction with the examples, but the embodiments of the present invention are not limited to the following examples.

[0052] The raw materials used in the present invention are all commercially available.

[0053] Example 1

[0054] (1) Using Zn-0.4Mn alloy (purchased from Ningbo Bowei Alloy Material Co., Ltd., wherein the manganese content is 0.40wt%) as a substrate, the substrate was polished in sequence using sandpaper with mesh numbers of 400#, 800#, 1200# and 2000# to obtain a pretreated substrate;

[0055] (2) placing the pretreated substrate obtained in step (1) into a reactor containing a calcium phosphate precursor solution (30 mL) of calcium nitrate (0.168 mol / L) and sodium dihydrogen phosphate (0.10 mol / L), the reaction temperature is 135° C., the reaction time is 5 h, and a calcium phosphate layer with a thickness of about 8 μm is loaded on the substrate; the surface of the calcium phosphate layer is laser remelted (the parameters of the laser remelting are start and end currents of 30 A; scanning speed of 200 mm / min; pulse frequency of 10 Hz; pulse duration of 1.5 ms; beam diameter of 0.8 mm; Ar gas protection flow rate of 20 L / min), to form a calcium phosphate-matrix mixed remelting zone, the thickness of the remelting zone is about 100-120 μm, and the surface structure of the medical degradable zinc-based material is obtained, and its surface and cross-sectional morphology are shown in FIG. Figure 1 As shown in (c) and (d) in .

[0056] Example 2

[0057] (1) Using Zn-0.35Li alloy (purchased from Ningbo Bowei Alloy Material Co., Ltd., wherein the lithium content is 0.35wt%) as a substrate, the substrate was polished in sequence using sandpaper with mesh sizes of 400#, 800#, 1200# and 2000# to obtain a pretreated substrate;

[0058] (2) placing the pretreated substrate obtained in step (1) into a reactor containing a calcium phosphate precursor solution (30 mL) of calcium nitrate (0.168 mol / L) and sodium dihydrogen phosphate (0.10 mol / L), the reaction temperature is 135° C., the reaction time is 5 h, and a calcium phosphate layer with a thickness of about 8 μm is loaded on the substrate; the surface of the calcium phosphate layer is laser remelted (the parameters of the laser remelting are: start and end current of 30 A; scanning speed of 200 mm / min; pulse frequency of 10 Hz; pulse duration of 2 ms; beam diameter of 0.8 mm; Ar gas protection flow rate of 20 L / min), to form a calcium phosphate-matrix mixed remelting zone, the thickness of the remelting zone is about 100-120 μm, and the surface structure of the medical degradable zinc-based material is obtained, and its cross-sectional metallographic image and cross-sectional morphology are shown in FIG. Figure 2 As shown in (b) and (d) in .

[0059] Comparative Example 1

[0060] The preparation method is the same as that of Example 1, except that in step (2), the pretreated substrate is directly laser remelted without loading the calcium phosphate layer to obtain the zinc-based material surface structure.

[0061] (1) Using Zn-0.4Mn alloy as a substrate, the substrate was polished in sequence using sandpapers with mesh sizes of 400#, 800#, 1200# and 2000# to obtain a pretreated substrate;

[0062] (2) The pretreated substrate obtained in step (1) was laser remelted (the parameters of laser remelting were: start and end current of 30 A; scanning speed of 200 mm / min; pulse frequency of 10 Hz; pulse duration of 1.5 ms; beam diameter of 0.8 mm; Ar gas protection flow rate of 20 L / min), and the surface structure of the zinc-based material after laser remelting was obtained. The surface and cross-sectional morphology are shown in FIG. Figure 1 As shown in (a) and (b) in .

[0063] Comparative Example 2

[0064] The preparation method is the same as that of Example 2, except that in step (2), the pretreated substrate is directly laser remelted without loading the calcium phosphate layer to obtain the zinc-based material surface structure.

[0065] (1) Using Zn-0.35Li alloy as a substrate, the substrate was polished in sequence using sandpaper with mesh sizes of 400#, 800#, 1200# and 2000# to obtain a pretreated substrate;

[0066] (2) The pretreated substrate obtained in step (1) was laser remelted (the parameters of laser remelting were: start and end current of 30 A; scanning speed of 200 mm / min; pulse frequency of 10 Hz; pulse duration of 2 ms; beam diameter of 0.8 mm; Ar gas protection flow rate of 20 L / min), and the surface structure of the zinc-based material after laser remelting was obtained. The cross-sectional metallographic image and cross-sectional morphology image are shown in FIG. Figure 2 As shown in (a) and (c) in .

[0067] Performance Testing

[0068] 1. Observation of surface structure and morphology of zinc-based materials

[0069] (1) Surface structure of zinc-based materials based on Zn-0.4Mn alloy

[0070] Figure 1 The surface and cross-sectional morphology of the surface structure of the zinc-based material obtained in Example 1 and Comparative Example 1 are as follows: Figure 1 As shown in (a) and (c), the surface of comparative example 1 is not loaded with a calcium phosphate layer, and after laser remelting, it presents a fish scale shape, while the surface of example 1 loaded with a calcium phosphate layer and then laser remelted presents a regularized surface, forming a dense protective layer on the surface of the zinc-based material; Figure 1 As shown in (b) and (d), the cross-sectional morphology shows that the laser melts the sample surface through energy transmission. The molten pool depth of Comparative Example 1 is about 120 μm. Due to the rapid cooling rate, the remelting zone presents dendritic grains and columnar grains with a growth direction perpendicular to the substrate surface. The grains in the remelting zone of Example 1 loaded with a calcium phosphate layer present a fine mesh.

[0071] (2) Surface structure of zinc-based materials based on Zn-0.35Li alloy

[0072] Figure 2 The cross-sectional metallographic images and cross-sectional morphology images of the surface structures of the zinc-based materials obtained in Example 2 and Comparative Example 2 are as follows: Figure 2 As shown in (a) and (b), there is a clear interface distinction between the remelting zone and the Zn-0.35Li alloy matrix in Comparative Example 1 and Example 1, and the remelting zone presents dendritic grain distribution. It can be observed that in the remelting zone of Example 1 after loading the calcium phosphate layer, the calcium phosphate layer is wrapped by the Zn-0.35Li alloy matrix after laser remelting, and the interface between the two is semicircular; Figure 2 As shown in (c) and (d), the cross-sectional morphology shows that the laser melts the sample surface through energy transmission. The molten pool depth of Comparative Example 2 is about 150 μm. Since Li is extremely active and has a low melting point, a deeper molten pool is formed and obvious columnar grains can be observed. In Example 2, a protective layer is formed on the surface of the Zn-0.35Li alloy matrix due to the loading of the calcium phosphate layer, which slows down heat conduction and thus causes a shallower molten pool depth.

[0073] 2. XRD test of surface structure of zinc-based materials

[0074] (1) Surface structure of zinc-based materials based on Zn-0.4Mn alloy

[0075] The surface structures of the Zn-0.4Mn alloy, the zinc-based materials obtained in Example 1 and Comparative Example 1 were analyzed by XRD spectra.

[0076] Figure 3 The XRD patterns of the surface structures of the zinc-based materials obtained from the Zn-0.4Mn alloy, Example 1 and Comparative Example 1 are as follows: Figure 3 As shown in (a), Zn-0.4Mn alloy is composed of Zn and MnZn 13 Phase composition, Zn-0.4Mn alloy (i.e., Zn-0.4Mn alloy, orange curve in the figure) shows stronger Zn peaks (i.e., 002, 100, 101, 102, 103, and 110 crystal planes) and weaker MnZn 13 Peaks (i.e. -331 and 221 crystal planes), while Comparative Example 1 (i.e. LSR-treated Zn-0.4Mn alloy, green curve in the figure) shows a strong Zn peak and a weak MnZn 13 Peak, Example 1 (i.e. LSR-treated Zn-0.4Mn alloy (with Ca 2+ , P 5 + ), the blue curve in the figure, also shows a strong Zn peak and a weak MnZn 13Peak. Figure 3 As shown in (b), Example 1 also shows CaZn2(PO4)2·2H2O peaks (i.e., -221, 112, 202 crystal planes) and Ca4H(PO4)3·2.5H2O peaks (i.e., -611, -2-12, 420, -322 crystal planes). The diffraction intensity of the Zn peak varies with the grain orientation, indicating that Zn 2+ Participates in the crystallization and formation of β-TCP phase during the hydrothermal process.

[0077] (2) Surface structure of zinc-based materials based on Zn-0.35Li alloy

[0078] The surface structures of the Zn-0.35Li alloy, the zinc-based materials obtained in Example 2 and Comparative Example 2 were analyzed by XRD spectra.

[0079] Figure 4 The XRD patterns of the surface structures of the zinc-based materials obtained from the Zn-0.35Li alloy, Example 2 and Comparative Example 2 are as follows: Figure 4 As shown in (a), the Zn-0.35Li alloy is composed of Zn and LiZn4 phases. The Zn-0.35Li alloy (i.e., Zn-0.35Lialloy, the orange curve in the figure) shows a strong Zn peak (i.e., 002, 100, 101, 102, 103 crystal planes) and a weak LiZn4 peak (i.e., 101 and 210 crystal planes). The Zn peak and LiZn4 peak in Comparative Example 1 (i.e., LSR-treated Zn-0.35Lialloy, the green curve in the figure) after laser remelting are both weakened. 2+ , P 5+ ), the blue curve in the figure, also shows a strong Zn peak and a weaker LiZn4 peak. Figure 4 As shown in (b), Example 1 also shows Ca9ZnLi(PO4)7 peaks (i.e., 125 and 0210 crystal planes), and the diffraction intensity of the Zn peak varies with the grain orientation, but the LiZn4 peak is also significantly enhanced, indicating that due to the hindrance of the calcium phosphate layer to heat conduction during the laser remelting process, the surface structure of the zinc-based material obtained in Example 1 cannot absorb and release heat normally, resulting in the precipitation of the LiZn4 phase.

[0080] 3. Surface hardness test

[0081] The Brinell hardness of the slices of the samples (Examples 1-2 and Comparative Examples 1-2) was tested using an in-situ nanoindenter (G200). Multiple point tests were performed from left to right along the direction perpendicular to the sample cross section. The first test point was about 40 μm from the sample surface, and the spacing between each test point was 50 μm. The test method was as follows: Figure 5As shown, the results are Figure 6 and Figure 8 shown.

[0082] The cross-sections of the samples (Examples 1-2 and Comparative Examples 1-2) were tested for microhardness using a hardness tester in a matrix-type dotting method (maintained for 15 seconds under a force of 25 g). The results are as follows: Figure 7 and Fig. 9 shown.

[0083] (1) Surface structure of zinc-based materials based on Zn-0.4Mn alloy

[0084] Figure 6 The surface strength test diagram of the zinc-based material surface structure prepared in Example 1 and Comparative Example 1 under the conditions of a load of 5 μN and a duration of 10 seconds is shown in FIG. Figure 6 As shown in (a) to (d) in the figure, the hardness of the Zn-0.4Mn alloy matrix is ​​0.6 to 0.8 GPa. After laser remelting, the hardness of the remelting zone of Comparative Example 1 and Example 1 is significantly improved to 1.5 to 2.0 GPa. This is because the MnZn is reduced after laser remelting. 13 Phase precipitation, some MnZn 13 The phase is solid dissolved into the matrix, making the microstructure of the material surface more uniform, and the lattice distortion caused by the solid-dissolved manganese atoms will enhance the dislocation slip resistance, thereby increasing the hardness of the surface structure of the zinc-based material.

[0085] Figure 7 The cross-sectional microhardness distribution diagram of the surface structure of the zinc-based material prepared in Example 1 and Comparative Example 1 is as follows: Figure 7 As shown in (a), after laser remelting, the comparative example 1 grows obvious columnar crystal regions from the Zn-0.4Mn alloy matrix. Figure 3 The XRD spectrum of Zn-0.4Mn alloy shows that there are more MnZn 13 Precipitated phase, after laser remelting, most of the MnZn 13 The precipitated phase is dissolved into Zn, which reduces the segregation of the structure and the brittleness of Zn-0.4Mn. From the boundary between the remelting zone and the Zn-0.4Mn alloy matrix, the hardness value of the remelting zone begins to increase. Figure 7 As shown in (b), after adding the calcium phosphate layer, most of the Ca and P melt into the sample surface, so the sample surface structure becomes rougher, and the sliding resistance between the grain boundaries on the sample surface is increased. The surface microhardness of Example 1 is improved. The experimental results show that the hardness of the Zn-0.4Mn alloy can be improved after adding a calcium phosphate layer and laser remelting.

[0086] (2) Surface structure of zinc-based materials based on Zn-0.35Li alloy

[0087] Figure 8 The surface strength test diagram of the zinc-based material surface structure prepared in Example 2 and Comparative Example 2 under the conditions of a load of 5 μN and a duration of 10 seconds is shown in FIG. Figure 8 As shown in (a) to (d), the hardness of the Zn-0.35Li alloy matrix is ​​1.20 to 1.30 GPa. After laser remelting, the hardness of the remelted zone of Comparative Example 2 and Example 2 is significantly improved, reaching 4.2 to 4.5 GPa. This is because the increase in Li solubility and the refinement of the microstructure after laser remelting reduce the eutectic structure of Zn+LiZn4. The increase in grain boundaries and phase boundaries will hinder the movement of dislocations, and the lattice distortion caused by the dissolved Li atoms will increase the slip resistance of dislocations, thereby increasing the hardness of the surface structure of the zinc-based material.

[0088] Fig. 9 The cross-sectional microhardness distribution diagram of the surface structure of the zinc-based material prepared in Example 2 and Comparative Example 2 is as follows: Fig. 9 As shown in (a), in Comparative Example 2 after laser remelting, the microhardness gradually increases from the substrate to the surface, and the hardness of the sample surface is about 1.3 to 1.6 times that of the Zn-0.35Li alloy matrix. In Example 2 loaded with a calcium phosphate layer, it can be observed that there is a clear dividing line between the remelting zone and the matrix, and the hardness increases from the intersection with the matrix to the sample surface.

[0089] 4. Friction and wear resistance test

[0090] The zinc-manganese alloy (Zn-0.4Mn alloy), zinc-lithium alloy (Zn-0.45Mn alloy), zinc-based material surface structures obtained in Examples 1-2 and Comparative Examples 1-2 were subjected to friction and wear tests using an Ampadon multifunctional high-temperature friction and wear tester, with the contact force set to 2N, the frequency set to 2Hz, and the friction time set to 1800s. After the friction test, the friction depth of the sample was measured using a surface profiler.

[0091] (1) Surface structure of zinc-based materials based on Zn-0.4Mn alloy

[0092] Table 1: Wear resistance test results of samples prepared in Example 1 and Comparative Example 1 and zinc-manganese alloy

[0093] sample Zinc-manganese alloy Comparative Example 1 Example 1 Friction coefficient 0.42 0.30 0.16 Wear depth(μm) 107.99 10.04 37.62

[0094] Fig.10 The wear resistance test diagram of the surface structure of the zinc-based material prepared by Zn-0.4Mn alloy, Example 1 and Comparative Example 1. Fig.10As shown in (a), after laser remelting, the roughness of the surface structure of the zinc-based material increases, and the contact area with the grinding ball increases, resulting in a decrease in its friction coefficient. According to the change in the friction coefficient, the curve of the Zn-0.4Mn alloy can be divided into two stages: in stage I, the friction coefficient fluctuates and rises with the sliding stroke; after entering stage II, the friction coefficient fluctuates the least and stabilizes near the average value. The two-stage curves of the sample in Example 1 are as follows: in stage I, as the sliding stroke progresses, the friction coefficient stabilizes near the average value; after entering stage II, the friction force fluctuates and rises, and gradually approaches the original friction coefficient of the Zn-0.4Mn alloy. The friction coefficient of the sample in comparative example 1 is stable near the average value, and the sample surface has stable wear resistance.

[0095] As shown in Table 1 and Fig.10 As shown in (b) to (d), the wear depth of the Zn-0.4Mn alloy is 107.99 μm, the wear depth of comparative example 1 is reduced to 10.04 μm, and the wear depth of the sample (Example 1) after loading the calcium phosphate layer and laser remelting is reduced to 37.62 μm.

[0096] (2) Surface structure of zinc-based materials based on Zn-0.35Li alloy

[0097] Table 2: Wear resistance test results of samples prepared in Example 2 and Comparative Example 2 and zinc-lithium alloy

[0098] sample Zinc-lithium alloy Comparative Example 2 Example 2 Friction coefficient 0.80 0.53 0.30 Wear depth(μm) 34.15 15.65 8.55

[0099] Fig.11 The wear resistance test diagram of the surface structure of the zinc-based material prepared by Zn-0.35Li alloy, Example 1 and Comparative Example 1. Fig.11 As shown in the figure, the friction coefficient of Zn-Li alloy is larger. Since the hardness and strength of Zn-Li itself are higher than those of zinc-manganese alloy, it exhibits better wear resistance and shallower wear scar depth. After laser remelting (Comparative Example 2), the wear resistance of the surface structure of the zinc-based material is significantly improved. If laser remelting is performed after loading the calcium phosphate layer (Example 2), with Ca 2+ , P 5+ With the addition of , wear resistance is significantly improved.

[0100] 5. Corrosion resistance test

[0101] The surface structures of zinc-based materials obtained from zinc-manganese alloy (Zn-0.4Mn alloy), zinc-lithium alloy (Zn-0.45Mn alloy), Examples 1-2 and Comparative Examples 1-2 were subjected to potentiodynamic polarization (PDP) and electrochemical impedance spectroscopy (EIS) tests by Autolab (PGSTAT 302) electrochemical workstation. The test used a three-electrode system, including a sample as a working electrode, a platinum sheet as a counter electrode, and a saturated calomel electrode as a reference electrode. The size of the test area was 8 mm × 8 mm, the test solution used was SBF, and the test was carried out at room temperature. The sample was coated with epoxy resin, exposing only the original sample and the laser remelted surface for comparison. Before the experiment, the sample was immersed in a 37°C SBF solution for 30 minutes to achieve a stable state and reduce errors.

[0102] (1) Surface structure of zinc-based materials based on Zn-0.4Mn alloy

[0103] Table 3: Corrosion resistance test results of samples prepared in Example 1 and Comparative Example 1 and zinc-manganese alloy

[0104] sample Zinc-manganese alloy Comparative Example 1 Example 1 <![CDATA[Corrosion current density (μA / cm 2 )]]> 23.74 16.52 24.23 Corrosion potential (V) -1.12 -1.12 -1.12 Corrosion rate (mm / year) 0.28 0.19 0.28

[0105] Fig.12 The corrosion resistance test diagram of the surface structure of the zinc-based material prepared by Zn-0.4Mn alloy, Example 1 and Comparative Example 1. Fig.12 As shown in (a), the Nyquist diagram shows that the samples before and after laser remelting show two capacitive arcs. The high-frequency capacitive arc represents the electron transfer process, and the low-frequency capacitive arc usually reflects the information of the corrosion product film on the metal surface. The laser remelting treated sample (Comparative Example 1) has a larger capacitive arc radius, indicating that the laser remelting treated sample has better corrosion resistance.

[0106] like Fig.12 As shown in the Bode curve of (b), there are two peaks in the Bode diagram, which matches the Nyquist diagram and the equivalent circuit of the double time constant R(Q(R(QR))). In this equivalent circuit, R s is the solution resistance, R cpf and Q cpf Represents the corrosion product layer resistance and capacitance, R ct and Q dl Represents the charge transfer resistance and capacitance between the sample and the solution. In the fitting process of the non-uniform interface reaction, the constant phase angle element Q is used to replace the ideal capacitor C. The impedance of Q is: ZQ = 1 / [Y(jw) n ], where Y is the value of capacitance and n is the exponent of Q. The corrosion product layer resistance R of the sample treated with LSR cpf (607.8Ω·cm -2) is greater than the R of the original sample cpf (202.9Ω·cm -2 ), indicating that the sample after laser remelting has better corrosion resistance, which is due to the MnZn 13 The phase was dissolved into the Zn matrix, causing micro-galvanic corrosion between the MnZn and the surface matrix Zn phase. 13 The amount of phase is reduced, thereby improving the corrosion resistance of the alloy.

[0107] like Fig.12 As shown in (c) in Figure 1, the polarization curves of the three samples before and after laser remelting are almost similar: the cathode shows an oxygen reduction reaction controlled by charge transfer, and the anode branch can clearly show the passivation-activation transition. However, compared with the original Zn-0.4Mn alloy, after direct laser remelting (Comparative Example 1), the anode reaction shifts to the left (reaction rate decreases) and the cathode reaction shifts to the right (reaction rate increases). 2+ and P 5+ In the laser remelted sample after the hydrothermal reaction (Example 1), the anode reaction shifted to the right (the reaction rate increased) while the cathode overlapped with the laser remelted sample.

[0108] As shown in Table 3, the corrosion potential and corrosion current density of Zn-0.4Mn alloy are -1.12 V and 23.74 μA / cm 2 The corrosion current density of comparative example 1 decreased to 16.52 μA / cm 2 The corrosion current density of Example 1 increased to 24.23 μA / cm 2 This shows that laser remelting can significantly enhance the corrosion resistance of Zn-0.4Mn alloy, but the corrosion resistance of laser remelting after covering the calcium phosphate layer is weakened, which is in line with the requirement of bone repair materials for certain corrosion resistance, making bone repair materials degradable and improving their biocompatibility.

[0109] (2) Surface structure of zinc-based materials based on Zn-0.35Li alloy

[0110] Table 4: Corrosion resistance test results of samples prepared in Example 2 and Comparative Example 2 and zinc-lithium alloy

[0111] sample Zinc-lithium alloy Comparative Example 2 Example 2 <![CDATA[Corrosion current density (μA / cm 2 )]]> 11.90 23.54 35.89 Corrosion potential (V) -1.17 -1.18 -1.19 Corrosion rate (mm / year) 0.20 0.27 0.39

[0112] Fig.13 The corrosion resistance test diagram of the surface structure of the zinc-based material prepared by Zn-0.35Li alloy, Example 1 and Comparative Example 1. Fig.13As shown in the figure, the polarization curve trends of the three samples before and after laser remelting are almost similar: the cathode shows an oxygen reduction reaction controlled by charge transfer, and the anode branch can clearly show the passivation-activation transition. However, compared with the original Zn-0.35Li alloy, the anode and cathode reaction rates of the laser remelted samples changed significantly. After LSR treatment, the anode branch changed slightly, while the cathode reaction shifted to the right (the reaction rate increased). This shows that laser remelting significantly reduced the corrosion resistance of the Zn-0.35Li alloy, while the loading of the calcium phosphate layer did not improve the corrosion resistance of the laser remelted surface. The Nyquist plot shows that the original sample has the largest capacitance arc radius, the capacitance arc radius of the sample after laser remelting is reduced, and the corrosion resistance is significantly reduced. The capacitance arc radius of the Zn-0.35Li alloy after laser remelting hydrothermal is the smallest, indicating that the corrosion resistance of the LSR surface is further reduced after the loading of the calcium phosphate layer. This is consistent with the requirement of certain corrosion resistance for bone repair materials, making the bone repair materials degradable and improving their biocompatibility.

[0113] 6. Biocompatibility testing

[0114] The extract was prepared according to ISO 10993-5:2009 standard, and the extract medium was cell culture medium (incomplete DMEM medium containing 12.5% ​​serum (containing double antibody)). The extraction ratio was the ratio of sample surface area to extract volume, which was 2:3. The extraction conditions were 37°C, 5% CO2 constant temperature incubator for 72 hours.

[0115] Preparation method of extract: a highly biocompatible zinc alloy surface structure is first ultrasonically cleaned with deionized water, and then moved into a biosafety cabinet after drying, and then immersed in 75% alcohol for 24 hours for disinfection, and then the zinc-manganese alloy is placed on filter paper for UV sterilization for 2 hours. After sterilization, the zinc alloy sample is placed in a centrifuge tube, a certain amount of cell culture medium is added, the centrifuge tube mouth is sealed, and placed in a 37°C, 5% CO2 incubator for 72 hours. 0.25mL of extract is extracted to prepare a 25% concentration of extract culture medium.

[0116] After L-929 cells (cell bank of the Chinese Academy of Sciences Typical Culture Collection Committee, numbered ***) were revived and passaged, 0.5% trypsin was used to detach the cells from the culture flask, and the detached cells were prepared into a cell suspension of 106 cells / mL using DMEM cell culture medium. Take a 96-well culture plate, add 100 μL of cell suspension to each well, and culture it in a 5% CO2 constant temperature incubator for 24 hours (37±2°C). After the cells adhere to the wall, remove the original culture medium and add 100 μL of extract to each well. After culturing in a 5% CO2 constant temperature incubator for 72 hours, use an inverted fluorescence microscope to obtain cell morphology.

[0117] (1) Surface structure of zinc-based materials based on Zn-0.4Mn alloy

[0118] Fig.14 The figure is a biocompatibility test diagram of the surface structure of zinc-based materials obtained by Zn-0.4Mn alloy, Example 1 and Comparative Example 1. After 72 hours of culture, the cell growth of the sample group treated with LSR is significantly better than that of the Zn-0.45Mn alloy group, and the cell growth advantage after loading the calcium phosphate layer is more obvious. For the Zn-0.45Mn alloy group, most of the cells in the extracted culture fluid died, showing obvious cytotoxicity.

[0119] Table 5: Fig.14 The content of elements in each group

[0120] sample Zn(mg / L) Mn(mg / L) Ca(mg / L) P (mg / L) Control group 0.02 ＜0.01 33.88 61.97 25% Zn-0.4Mn alloy leaching solution group 5.81 0.01 33.39 64.79 25% Comparative Example 1 Extract Group 8.39 0.03 36.96 63.67 25% Example 1 extract group 6.73 0.03 38.05 69.92

[0121] As shown in Table 5, in Example 1, the calcium phosphate layer is loaded and then laser remelted, and the precipitated Ca 2+ and P 5+ It promoted the growth of L-929 fibroblasts and improved their cell compatibility.

[0122] (2) Surface structure of zinc-based materials based on Zn-0.35Li alloy

[0123] Fig.15 Figure 2 is a biocompatibility test diagram of the surface structure of zinc-based materials prepared by Zn-0.35Li alloy, Example 2 and Comparative Example 2. In the sample group cultured for 24 hours, due to the presence of Zn+LiZn4 eutectic structure in the original sample, the electrode potential E° (-3.04 V) of Li is much lower than that of Zn (-0.76 V), forming a potential difference that accelerates corrosion and leads to the precipitation of Zn 2+ Also more, and too high concentration of Zn 2+ The cells in the Zn-0.35Li alloy group showed cytotoxicity, so the number of cells attached was the least. After 72 hours of culture, the cell growth of Comparative Example 2 was almost the same as that of the Zn-0.35Li alloy group, and the cell growth of Example 2 was relatively stable. This indicates that after the calcium phosphate layer was covered in Example 2, the precipitated Ca 2+ and P 5+ It promoted the growth of L-929 fibroblasts and improved their cell compatibility.

[0124] The above description is only a preferred embodiment of the present invention and is not intended to limit the present invention. Although the present invention is described in detail with reference to the aforementioned embodiments, those skilled in the art can still modify the technical solutions described in the aforementioned embodiments or replace some of the technical features therein by equivalents. Any modification, equivalent replacement, improvement, etc. made within the spirit and principle of the present invention shall be included in the protection scope of the present invention.

Claims

1. A medical degradable zinc-based material surface structure, comprising a zinc-based matrix, characterized in that: A calcium phosphate layer is loaded on the zinc-based substrate, and the surface of the calcium phosphate layer is laser remelted to form a remelted zone.

2. The medical degradable zinc-based material surface structure according to claim 1, characterized in that: The zinc-based matrix is ​​zinc or a zinc alloy, and the zinc alloy is an alloy formed by zinc and at least one metal selected from manganese, iron, lithium, magnesium, copper and calcium; in the zinc alloy, 50wt%≤Zn<100wt%.

3. The medical degradable zinc-based material surface structure according to claim 2, characterized in that: The zinc alloy is a zinc-manganese alloy or a zinc-lithium alloy, and the content of manganese or lithium in the zinc alloy is 0.30-0.45wt%.

4. The medical degradable zinc-based material surface structure according to claim 1, characterized in that: In the surface structure of the zinc-based material, 0<the thickness of the calcium phosphate salt layer is ≤20μm, and 0<the thickness of the remelting zone is ≤300μm.

5. The method for preparing the medical degradable zinc-based material surface structure according to any one of claims 1 to 4, characterized in that: The following steps are involved: (1) removing the oxide layer from the surface of the substrate to obtain a pretreated substrate; (2) A calcium phosphate layer is loaded on the pretreated substrate obtained in step (1), and the surface of the calcium phosphate layer is laser remelted to obtain a medical degradable zinc-based material surface structure.

6. The method for preparing the medical degradable zinc-based material surface structure according to claim 5, characterized in that: In step (2), the method for loading the calcium phosphate layer is to place the pretreated substrate in a calcium phosphate precursor solution for a hydrothermal reaction.

7. The method for preparing the medical degradable zinc-based material surface structure according to claim 6, characterized in that: The calcium-phosphorus precursor solution consists of calcium nitrate and sodium dihydrogen phosphate.

8. The method for preparing the medical degradable zinc-based material surface structure according to claim 7, characterized in that: In the calcium-phosphorus precursor solution, the molar ratio of calcium nitrate to sodium dihydrogen phosphate is 1-2:

1.

9. The method for preparing the medical degradable zinc-based material surface structure according to claim 5, characterized in that: In step (2), the pulse frequency of the laser remelting is 5 to 10 Hz; and the pulse duration is 1 to 2 ms.

10. Use of the medical degradable zinc-based material surface structure according to any one of claims 1 to 4 in the preparation of bone healing materials.

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