A non-homogenous metal bone implant, method of making and use thereof

Abstract

This invention relates to the field of biomedical materials technology, specifically to a method and application for the fabrication of heterogeneous metal bone implants. The fabrication method includes the following steps: S1: obtaining a physical model of the bone implant; S2: generating an L-PBF printing file based on the physical model of the bone implant to obtain a porous model of the bone implant; S3: based on the L-PBF printing file, using multi-material laser powder bed melting technology, performing integrated multi-material L-PBF printing with two or more different metal powders having large differences in melting and boiling points to obtain a porous heterogeneous bone implant blank; finally, post-processing the porous heterogeneous bone implant blank to obtain a heterogeneous metal bone implant. The fabrication method of this invention solves the core technical problem of weak interfacial bonding and easy failure in multi-material implants, and also addresses the problem of large differences in melting and boiling points between two metal materials.

Description

Technical Field

[0001] This invention relates to the field of biomedical materials technology, specifically to a method for preparing and applying a heterogeneous metal bone implant. Background Technology

[0002] Bone defects are one of the major challenges facing the global orthopedic field, with over 2 million cases annually. Currently, metallic implants (such as stainless steel and titanium alloys) have become the mainstream choice for bone repair surgery due to their excellent mechanical strength and good biocompatibility. However, traditional single-material systems have inherent limitations: the high elastic modulus of inert metals can lead to stress shielding, accelerating bone resorption, and also requiring secondary surgery for removal, while biodegradable metals face the risk of degradation behavior not matching the rate of bone regeneration. To overcome the limitations of single materials, composite material bone implants have become a research hotspot.

[0003] However, traditional manufacturing processes (such as casting, forging, and machining) struggle to achieve precise control over the spatial distribution of multiple materials and to form complex internal porous structures, failing to meet personalized clinical needs. Although laser powder bed fusion (L-PBF) technology offers a new approach to manufacturing customized porous structures, current research remains focused on single materials or simple functionally graded materials (such as the pore gradient variation of titanium alloys), with almost no exploration of multi-material porous metal implants. Therefore, developing a new generation of implants that combine multifunctionality and customization has become a pressing technical challenge that needs to be addressed. Summary of the Invention

[0004] To address the aforementioned technical problems, this invention provides a heterogeneous metal bone implant, its preparation method, and its application. This invention selects two biocompatible metal materials with different electrode potentials and employs multi-material laser powder bed melting (L-PBF) technology. Through dynamic powder switching and interface remelting processes, metallurgical bonding is achieved at the interface of the heterogeneous materials, significantly improving the interfacial bonding strength and fracture resistance. This ensures that the heterogeneous implant with a complex three-dimensional interconnected pore structure maintains structural integrity and functionality under harsh physiological conditions in vivo. It solves the core technical problem of weak interfacial bonding and easy failure in multi-material implants and addresses the significant difference in melting and boiling points between the two metal materials.

[0005] Therefore, the present invention provides the following technical solution:

[0006] In a first aspect, the present invention provides, in an optional embodiment, a method for preparing a heterogeneous metal bone implant, comprising the following steps:

[0007] S1: Based on the three-dimensional image data of the bone defect site, reconstruct the bone defect model, and define the spatial distribution and geometric configuration of two or more different metal materials with large melting and boiling point differences in the bone defect model to obtain the solid model of the bone implant.

[0008] S2: Based on the physical model of the bone implant, pore structure parameters are independently designed for the two different metal material regions, and an L-PBF printing file is generated to obtain a porous model of the bone implant.

[0009] S3: Under the protection of inert gas, based on the L-PBF printing file, multi-material laser powder bed melting technology is used to perform multi-material L-PBF integrated printing using two or more different metal material powders with large melting and boiling point differences to obtain a porous heterogeneous bone implant blank. Finally, the porous heterogeneous bone implant blank is post-processed to obtain a heterogeneous metal bone implant.

[0010] Among them, the two or more different metal materials with large melting and boiling point differences are selected from metal systems with biocompatibility and different electrode potentials. By forming an electrochemical dipole in the body fluid environment, microcurrents are generated to simulate endogenous bioelectric signals to accelerate bone integration and regulate the degradation behavior of different materials.

[0011] In this invention, two biocompatible metal materials with different electrode potentials form an electrochemical dipole in a body fluid environment to generate a microcurrent that simulates endogenous bioelectric signals to accelerate bone integration and regulate the degradation behavior of different materials.

[0012] Furthermore, in step S1, the two different metallic materials are Ti-based metals and Zn-based metals;

[0013] The Ti-based metal is selected from one or more of pure Ti, Ti6Al4V alloy, or Ti-0.3Mo-0.8Ni alloy, and the Zn-based metal is selected from one or more of pure Zn, Zn-Mg alloy, or Zn-Li alloy.

[0014] In this invention, the regulation of the electrochemical dipole aims to achieve temporal control and functionalized output of degradation behavior: the metal, acting as the anode, degrades preferentially, and its degradation rate can be regulated through material ratios, structural design, and interface geometry; the metal, acting as the cathode, is protected from degradation to provide long-term stable mechanical support; the degradation behavior is associated with the following biological functions:

[0015] (a) When Zn-based metals are used as anodes for degradation, the Zn released during degradation 2+ At effective concentrations, the ions exhibit both antibacterial and osteogenic activities.

[0016] (b) When Ti-based metals are used as cathodes, their degradation is delayed, providing long-term mechanical function for the main load-bearing areas.

[0017] Furthermore, the degradation rate is actively regulated through one or more of the following design techniques:

[0018] (1) Composition and spatial proportion design: control the relative volume fraction of different metal materials;

[0019] (2) Structural design: Independently design the macroscopic structure and pore units of each metal region to change its specific surface area;

[0020] (3) Interface design: Increase the number of interface or electrochemical dipole repeating units.

[0021] Furthermore, for combinations containing Ti-based metals, their behavior as cathodes in actual physiological environments stems from the alteration of their electrochemical activity by a stable oxide film formed in situ on their surface.

[0022] Preferably, in step S3, during the integrated printing of the multi-material L-PBF, after switching from one metal powder to another, a layer of metal powder to be printed is laid on the surface of the previously printed metal powder layer, and an interface remelting process is performed according to the laser parameters of the layer to be printed. When the metal powder is Ti-based, the laser parameters of the laser powder bed are: power 120-300W, scanning speed 200-600mm / s, layer thickness 20-50μm; and / or, when the metal powder is Zn-based, the laser parameters of the laser bed are: power 30-120W, scanning speed 100-800mm / s, layer thickness 20-40μm. The metal powder is spherical or nearly spherical, with a particle size distribution D. 50 The particle size D of the different metal powders used in the same printing process is 10-60 μm. 50 The difference is no more than 30μm.

[0023] Preferably, in step S2, the pore structure parameters include pore unit type, pore unit size, rod diameter, and unit structure porosity; the pore unit type includes one or a combination of structures based on triple periodic minimal surfaces, lattice structures, anisotropic structures, functionally graded porous structures, and random biomimetic structures; the pore unit size is 1-5 mm; the rod diameter is 0.4-4 mm; and the unit structure porosity is 40-95%.

[0024] Furthermore, when the porosity difference of adjacent metal materials in unit structures is not less than 10%, a continuous gradient variation region of the porosity of the unit structure is set, and the porosity variation gradient of the unit structure is not greater than 4% / mm.

[0025] Preferably, in step S1, the spatial distribution and several configurations of the two different metal materials defined in the bone defect model are defined based on the biomechanical environment, degradation sequence and functional requirements.

[0026] Preferably, in step S3, the inert gas is selected from one or more of nitrogen, argon, or neon; and / or, the post-processing includes wire cutting to remove the support structure, cleaning, sandblasting, abrasive flow polishing, heat treatment, and surface functional treatment.

[0027] Secondly, the present invention provides a heterogeneous metal bone implant obtained by the above preparation method.

[0028] Preferably, the heterogeneous metal bone implant has a porous structure with a three-dimensional interconnected pore network, and the pore structure parameters in different metal material regions can be independently designed and controlled; in the heterogeneous metal bone implant, the two different metal materials are directly bonded at the interface without gradient component transition.

[0029] Preferably, the heterogeneous metal bone implant has a layered configuration of at least two layers, and the shape of the layers can be biomimetically designed according to the skeletal anatomy of the implantation site. Furthermore, the layered configuration can be two, three, or four layers, but is not limited to these numbers; it can be a layered configuration with more than four layers. By adjusting the composition, porosity, or microstructure of the material layer by layer, the layered configuration achieves a gradient transition between mechanical properties and biological functions on a macroscopic level, thereby better matching the biological and mechanical environment of the host bone.

[0030] Thirdly, the present invention provides an application of the above-mentioned heterogeneous metal bone implant in the preparation of orthopedic and / or dental metal implants.

[0031] Compared with the prior art, the present invention has one of the following beneficial effects:

[0032] 1. This invention selects two biocompatible metal materials with different electrode potentials and employs multi-material laser powder bed melting (L-PBF) technology. Through dynamic powder switching and interface remelting processes, metallurgical bonding is achieved at the interface of heterogeneous materials, significantly improving the interface bonding strength and fracture resistance. This ensures that heterogeneous implants with complex three-dimensional through-pore structures maintain structural integrity and functionality under harsh physiological conditions in vivo. It solves the core technical problem of weak interface bonding and easy failure in multi-material implants and addresses the issue of significant differences in melting and boiling points between the two metal materials.

[0033] 2. This invention effectively eliminates the need for secondary surgery through the component partitioning design of biodegradable materials (such as Zn-based metals) and high-strength materials (such as Ti-based metals); the application of antibacterial and osteogenic functional materials significantly reduces the risk of infection and accelerates bone tissue regeneration, achieving a breakthrough in clinical treatment.

[0034] 3. This invention achieves temporal control over the degradation behavior, antibacterial effect, and osteogenic activity of implants in vivo by spatially partitioning biometals with different electrode potentials (such as Ti-based and Zn-based) and synergistically regulating their pore structure and interface geometry. The implant can actively release bioactive ions (such as Zn) in the early stages. 2+ It has antibacterial and bone-accelerating properties, and gradually degrades in the later stages. At the same time, it is supported by high-strength material regions (such as Ti-based) for long-term mechanical support, thereby eliminating the need for secondary surgery and realizing a functional leap from "passive support" to "active repair".

[0035] 4. By independently designing pore structure parameters (such as porosity and rod diameter) for different material regions, this invention allows for gradient control of the overall stiffness and elastic modulus of the implant, effectively avoiding stress shielding effects. Its interconnected porous structure across materials not only provides space for cell migration and tissue ingrowth but also ensures efficient nutrient transport, ultimately achieving integrated optimization of the implant's mechanical properties and biological functions, greatly promoting bone integration efficiency. Attached Figure Description

[0036] To more clearly illustrate the technical solutions in the embodiments of the present invention, the accompanying drawings used in the embodiments will be briefly introduced below. Obviously, the drawings described below are only some embodiments of the present invention. For those skilled in the art, other drawings can be obtained based on these drawings without creative effort.

[0037] Figure 1 This is a schematic diagram of the preparation process of the heterogeneous metal bone implant provided in Embodiment 2 of the present invention;

[0038] Figure 2 The diagrams and general layouts of 1-Zn / Ti designed in Embodiment 1 and 3-Zn / Ti designed in Embodiment 2 of the present invention are shown, wherein a is 1-Zn / Ti and b is 3-Zn / Ti.

[0039] Figure 3 The preparation and characterization of the titanium-zinc composite material provided in Comparative Example 2 of the present invention are shown in Figure a, where a is a schematic diagram of the preparation process and b is a multi-scale microstructure analysis of the sintered Ti-Zn sample.

[0040] Figure 4Microscopic characterization images of the heterogeneous metal bone implant 1-Zn / Ti prepared in Example 1, the heterogeneous metal bone implant 3-Zn / Ti prepared in Example 2, and the pure Zn metal bone implant prepared in Comparative Example 1 are shown. Among them, a is the microscopic characterization image of pure Zn, b is the microscopic characterization image of 1-Zn / Ti, c is the microscopic characterization image of 3-Zn / Ti, and d is a schematic diagram of the interface bonding of Zn / Ti.

[0041] Figure 5 Compression strength diagrams of heterogeneous metal bone implant 1-Zn / Ti prepared in Example 1, heterogeneous metal bone implant 3-Zn / Ti prepared in Example 2, and pure Zn metal bone implant prepared in Comparative Example 1;

[0042] Figure 6 Degradation behavior diagrams of heterogeneous metal bone implant 1-Zn / Ti prepared in Example 1, heterogeneous metal bone implant 3-Zn / Ti prepared in Example 2, and pure Zn metal bone implant prepared in Comparative Example 1;

[0043] Figure 7 In vitro cell compatibility diagrams of heterogeneous metal bone implant 1-Zn / Ti prepared in Example 1, heterogeneous metal bone implant 3-Zn / Ti prepared in Example 2, and pure Zn metal bone implant prepared in Comparative Example 1;

[0044] Figure 8 The images show in vivo osteogenic and degradation patterns of the heterogeneous metal bone implant 1-Zn / Ti prepared in Example 1, the heterogeneous metal bone implant 3-Zn / Ti prepared in Example 2, and the pure Zn metal bone implant prepared in Comparative Example 1. In these images, a is a Micro-CT scan result, b is a quantitative analysis of osteogenic indices, c is a hard histological section of the tissue, and d is a SEM image and EDS mapping of the hard tissue cross-section (Ti: Ti matrix, Zn: Zn matrix, NB: new bone, DP: degradation products).

[0045] Figure 9 This is an ion release diagram of the heterogeneous metal bone implant Mg / Zn prepared in Example 3 of the present invention, where a is Mg 2+ Ion concentration, where b is the Zn ion concentration;

[0046] Figure 10 The degradation behavior of the titanium-zinc composite material provided in Comparative Example 2 of the present invention is shown in the figure. In the figure, a represents the degradation form of Ti-20, Ti-30Zn and Ti-90Zn after different immersion durations, b represents EDS mapping, c represents FTIR spectrum, d represents XRD pattern of Ti-90Zn degradation product, e represents Zn ion concentration in SBF solution, f represents weight loss, and g represents comparison of degradation rate of Ti-Zn sample with other biodegradable metals.

[0047] Figure 11 The in vitro biocompatibility of the titanium-zinc composite material provided in Comparative Example 2 of the present invention is shown in the figure, where a is the cytotoxicity CCK-8 result and b is the live / dead cell staining and cytoskeleton staining.

[0048] Figure 12 The in vivo osteogenic behavior of the titanium-zinc composite material provided in Comparative Example 2 of this invention is shown in the following figures: a) is a flowchart of the animal experiment; b) is a Micro-CT scan result; c) is a quantitative analysis of osteogenic indices; and d) is Masson staining of tissue sections (EB: bone present, NB: new bone, FT: fibrous tissue, DP: degradation products).

[0049] Figure 13 SEM images and EDS mappings of the hard tissue cross sections of Ti-30Zn and Ti-90Zn groups in the titanium-zinc composite material provided in Comparative Example 2 of this invention (EB: existing bone, NB: new bone, FT: fibrous tissue, DP: degradation products).

[0050] Figure 14 This is a microstructure diagram of the Zn / Ti composite metal bone implant prepared in Comparative Example 3 of the present invention. Detailed Implementation

[0051] To make the objectives, technical solutions, and advantages of this invention clearer, the invention will be further described in detail below with reference to specific embodiments. It should be understood that the specific embodiments described herein are merely illustrative of the invention and are not intended to limit the invention.

[0052] The following examples and comparative examples all use heterogeneous metal bone implants as examples to demonstrate the preparation of porous metal composite materials.

[0053] The experimental materials and equipment involved in the examples mainly include, but are not limited to:

[0054] Reagents:

[0055] Table 1 Main Reagents

[0056]

[0057] instrument:

[0058] Table 2 Main Instruments

[0059]

[0060] The technical solution of the present invention will be described below with reference to the embodiments.

[0061] Example 1

[0062] This embodiment provides a method for preparing a porous titanium / zinc heterogeneous metal bone implant with a stacked configuration (number of stacks = 2), including the following steps:

[0063] Step 1: 3D Modeling and Composition Partitioning

[0064] Based on CT data of femoral condyle bone defects in rats, a three-dimensional model was reconstructed using software. The implant size was preset to φ2.8mm × 4.8mm, and its axis was divided into two zones, each containing one of two metals. Titanium (Ti) was selected as a bone repair material due to its excellent biocompatibility and mechanical properties among biomedical metals. However, considering its non-degradability, insufficient antibacterial properties, and stress shielding risks, a biodegradable metal was incorporated into the design. This metal can gradually degrade in body fluids and promote tissue repair through its degradation products. Zn-based metals possess biodegradability, antibacterial properties, and the ability to promote bone regeneration. Pure Zn has poor mechanical properties (yield strength of 30MPa and elongation of 2%), making it unsuitable for load-bearing bone defects. Furthermore, the degradation rate of Zn-based metals is slower than the bone regeneration rate, requiring acceleration. Based on this, the present invention selects pure Ti and pure Zn to be deployed in two partitions respectively, wherein pure Ti region 5 provides mechanical support, and pure Zn region 6 degrades and has both antibacterial and osteogenic functions. Because the material has a Ti-Zn interface, it is named 1-Zn / Ti.

[0065] Step 2: Design of cross-material pore structure

[0066] This invention is based on the triple-period minimum surface (TPMS) with its zero average cell curvature and large surface area. The TPMS-based scaffold has a smooth transition, enabling uniform stress distribution. It also possesses topological features and curvature similar to human bone trabeculae, thus improving tissue regeneration. Furthermore, considering the significant impact of TPMS-based scaffold pore size on biocompatibility and osteogenic activity, this invention selected a pore size of 400–600 μm as suitable for osteogenic and osteointegrative processes, and found that higher porosity leads to faster material degradation. Therefore, based on the invention's objective of accelerating degradation and promoting osteogenic growth, the pore structure chosen is the Gyriod structure from TPMS, with a porosity of 60%.

[0067] Pure Ti region 5: pore unit size 2mm, porosity 60%, adopting the Gyriod structure in TPMS.

[0068] Pure Zn region 6: pore unit size 2mm, porosity 60%, using the Gyriod structure in TPMS.

[0069] Interface gradient transition layer: Since there is no difference in porosity, no transition layer is set.

[0070] Step 3: Multi-material L-PBF printing

[0071] Pure Ti and pure Zn powders with a particle size distribution of 15~53μm were prepared by vacuum inert gas atomization. Before printing, the powders were dried at 100℃ for 5h to remove moisture. The oxygen content was controlled to ≤100ppm using an L-PBF machine under argon protection.

[0072] Laser parameters:

[0073] Pure Ti region: Laser 1 power 80W, scanning speed 450mm / s, incubation space 80μm, layer thickness 40μm.

[0074] Pure Zn region: Laser power 60W, scanning speed 350mm / s, incubation space 70μm, layer thickness 30μm.

[0075] Since Zn has a lower melting and boiling point than Ti (420℃ and 907℃ respectively), Zn evaporation during material preparation can lead to severe interfacial cracks or porosity. To address this issue, this invention employs an interfacial remelting process, which involves performing two laser scans on the Ti-Zn interface using the parameters of the material to be printed.

[0076] Material replacement process: After completing the pure Ti layer printing, replace the powder feeder with pure Zn powder 4, clean the residual pure Ti powder 2 from the forming chamber, adjust the laser parameters, and complete the material printing (see...). Figure 2 a).

[0077] Step 4: Post-processing

[0078] After the printed parts cool to room temperature, the powder in the molding chamber is recovered, the parts are separated from the substrate by wire cutting, the powder on the surface is cleaned by compressed air, and vacuum annealing is performed to optimize the mechanical properties.

[0079] Example 2

[0080] See Figure 1 This embodiment provides a method for preparing a laminated (layer number = 4) titanium / zinc bimetallic bone implant, including the following steps:

[0081] Step 1: 3D Modeling and Composition Partitioning

[0082] Based on CT data of rat femoral condyle defects, a three-dimensional model was reconstructed using software. The implant size was preset to φ2.8mm × 4.8mm, and its axis was divided into four zones, with Zn and Ti metals alternately deployed in sequence. Because the material has three Ti-Zn interfaces, it was named 3-Zn / Ti.

[0083] The pure Ti region 5 provides mechanical support, while the pure Zn region 6 degrades and has both antibacterial and osteogenic functions.

[0084] Step 2: Design of cross-material pore structure

[0085] Pure Ti region 5: pore unit size 2mm, porosity 60%, adopting the Gyriod structure in TPMS.

[0086] Pure Zn region 6: pore unit size 2mm, porosity 60%, using the Gyriod structure in TPMS.

[0087] Interface gradient transition layer: Since there is no difference in porosity, no transition layer is set.

[0088] Step 3: Multi-material L-PBF printing

[0089] Pure Ti and pure Zn powders with a particle size distribution of 15–53 μm were prepared by vacuum inert gas atomization. Before printing, the powders were dried at 100 °C for 5 h to remove moisture. The oxygen content was controlled to ≤100 ppm using an L-PBF machine under argon protection.

[0090] Laser parameters:

[0091] Pure Ti region: Laser 1 power 80W, scanning speed 450mm / s, incubation space 80μm, layer thickness 40μm.

[0092] Pure Zn region: Laser power 60W, scanning speed 350mm / s, incubation space 70μm, layer thickness 30μm.

[0093] Interface remelting: Each interface zone is remelted using the laser parameters of the material to be printed.

[0094] Material replacement procedure: Please refer to Figure 1 After completing the pure Ti layer printing, replace the powder feeder with pure Zn powder 4, clean the residual pure Ti powder 2 from the forming chamber, adjust the laser parameters, and continue printing. Repeat this alternating process to complete the bracket printing. Please refer to [link to documentation]. Figure 2 b.

[0095] Step 4: Post-processing

[0096] After the printed parts cool to room temperature, the powder in the molding chamber is recovered, the parts are separated from the substrate by wire cutting, the powder on the surface is cleaned by compressed air, and vacuum annealing is performed to optimize the mechanical properties.

[0097] Example 3

[0098] See Figure 1 This embodiment provides a method for preparing a laminated (layer number = 3) titanium / zinc bimetallic bone implant, including the following steps:

[0099] Step 1: 3D Modeling and Composition Partitioning

[0100] Based on CT data of rat femoral condyle defects, a three-dimensional model was reconstructed using software. The implant size was preset to φ2.8mm × 4.8mm, and its axis was divided into three partitions, with Zn and Ti metals alternately deployed in sequence. Because the material has three Ti-Zn interfaces, it was named 2-Zn / Ti.

[0101] The pure Ti region 5 provides mechanical support, while the pure Zn region 6 degrades and has both antibacterial and osteogenic functions.

[0102] Step 2: Design of cross-material pore structure

[0103] Pure Ti region 5: pore unit size 2mm, porosity 60%, adopting the Gyriod structure in TPMS.

[0104] Pure Zn region 6: pore unit size 2mm, porosity 60%, using the Gyriod structure in TPMS.

[0105] Interface gradient transition layer: Since there is no difference in porosity, no transition layer is set.

[0106] Step 3: Multi-material L-PBF printing

[0107] Pure Ti and pure Zn powders with a particle size distribution of 15–53 μm were prepared by vacuum inert gas atomization. Before printing, the powders were dried at 100 °C for 5 h to remove moisture. The oxygen content was controlled to ≤100 ppm using an L-PBF machine under argon protection.

[0108] Laser parameters:

[0109] Pure Ti region: Laser 1 power 80W, scanning speed 450mm / s, incubation space 80μm, layer thickness 40μm.

[0110] Pure Zn region: Laser power 60W, scanning speed 350mm / s, incubation space 70μm, layer thickness 30μm.

[0111] Interface remelting: Each interface zone is remelted using the laser parameters of the material to be printed.

[0112] Material replacement procedure: Please refer to Figure 1 After completing the pure Ti layer printing, the powder feeding chamber was replaced with pure Zn powder 4, the residual pure Ti powder 2 in the forming chamber was cleaned, the laser parameters were adjusted and printing continued, alternating and repeating until the bracket printing was completed.

[0113] Step 4: Post-processing

[0114] After the printed parts cool to room temperature, the powder in the molding chamber is recovered, the parts are separated from the substrate by wire cutting, the powder on the surface is cleaned by compressed air, and vacuum annealing is performed to optimize the mechanical properties.

[0115] Comparative Example 1

[0116] This comparative example provides a method for preparing a pure Zn metal bone implant, including the following steps:

[0117] Step 1: 3D Modeling

[0118] Based on CT data of femoral condyle defects in rats, a three-dimensional model was reconstructed using software, with the implant size preset to φ2.8mm×4.8mm.

[0119] Step 2: Pore Structure Design

[0120] The pore unit size is 2mm, the porosity is 60%, and the Gyriod structure in TPMS is adopted.

[0121] Step 3: L-PBF Printing

[0122] Pure Zn powder with a particle size distribution of 15-53 μm was prepared by vacuum inert gas atomization. Before printing, the powder was dried at 100℃ for 5 hours to remove moisture. Printing was completed using an L-PBF machine under argon protection, with the oxygen content controlled to ≤100 ppm.

[0123] Laser parameters: laser power 60W, scanning speed 350mm / s, incubation space 70μm, layer thickness 30μm.

[0124] Step 4: Post-processing

[0125] After the printed parts cool to room temperature, the powder in the molding chamber is recovered, the parts are separated from the substrate by wire cutting, the powder on the surface is cleaned by compressed air, and vacuum annealing is performed to optimize the mechanical properties.

[0126] Comparative Example 2

[0127] This comparative example provides a method for preparing a Ti / Zn composite metal bone implant, including the following steps:

[0128] Step 1: Size Preset

[0129] Cylindrical Ti-Zn samples with a diameter of 20 mm × 13 mm were used for in vitro experiments, while cylindrical Ti-Zn samples with a diameter of 2.5 mm × 4 mm were used for in vivo implantation.

[0130] Step 2: Powder Mixing

[0131] High-purity Ti (TA1, 50–100 μm, Peshing New Metal, Changzhou, China) and Zn (99.99%, 5–15 μm, China New Metal, Beijing, China) powders were uniformly mixed at different mass ratios (wt%) for 7 h using a planetary ball mill at 250 rpm.

[0132] Step 3: Powder Sintering

[0133] See Figure 3 a. A designed SPS process was implemented on titanium-based and zinc-based systems in a vacuum sintering furnace (SPS 320MKII, Sumitomo, Japan). The SPS process for titanium-based composites (Zn ≤ 50 wt%) consisted of two steps: sintering at 400℃ and 10 MPa for 15 minutes, followed by sintering at 700℃ and 50 MPa for 10 minutes. The Zn-based composites (Zn ≥ 50 wt%) were sintered using a single-step process: sintering at 350℃ and 50 MPa for 25 minutes. The SPS heating rate was 100℃ / min. The composites were named Ti-xZn (x = 3, 5, 7, 10, 20, 30, ..., 90, in weight percentage).

[0134] Comparative Example 3

[0135] This comparative example provides a method for preparing a Zn / Ti composite metal bone implant, including the following steps:

[0136] Step 1: Size Preset

[0137] The preset size is φ2.8mm×4.8mm. Its axis is divided into two sections, and two types of metals are deployed in each section.

[0138] Step 2: Design of cross-material pore structure

[0139] Pure Ti region: pore unit size 2mm, porosity 60%, using the Gyriod structure in TPMS.

[0140] Pure Zn region: pore unit size 2mm, porosity 60%, adopts the Gyriod structure in TPMS, without Zn-Ti transition layer.

[0141] Step 3: Multi-material L-PBF printing

[0142] Pure Ti and pure Zn powders with a particle size distribution of 15~53μm were prepared by vacuum inert gas atomization. Before printing, the powders were dried at 100℃ for 5h to remove moisture. The oxygen content was controlled to ≤100ppm using an L-PBF machine under argon protection.

[0143] Laser parameters:

[0144] Pure Ti region: laser power 80W, scanning speed 450mm / s, incubation space 80μm, layer thickness 40μm.

[0145] Pure Zn region: laser power 60W, scanning speed 350mm / s, incubation space 70μm, layer thickness 30μm.

[0146] The Zn-Ti interface remelting process is not performed.

[0147] Material replacement process: After completing the pure Ti layer printing, replace the powder feeder with pure Zn powder, clean the residual pure Ti powder from the forming chamber, adjust the laser parameters, and then complete the material printing (see...). Figure 2 a).

[0148] Step 4: Post-processing

[0149] After the printed parts cool to room temperature, the powder in the molding chamber is recovered, the parts are separated from the substrate by wire cutting, the powder on the surface is cleaned by compressed air, and vacuum annealing is performed to optimize the mechanical properties.

[0150] Comparative Example 4

[0151] This comparative example provides a method for preparing a Zn / Ti composite metal bone implant, including the following steps:

[0152] Sample preparation: Using the same L-PBF process parameters as in Example 1, porous composite samples with Ti-Zn interfaces were prepared.

[0153] Remelting process: The printed sample is placed in a traditional box-type heat treatment furnace protected by argon gas.

[0154] Remelting temperature: Two temperature points were set: 420℃ (slightly higher than the melting point of Zn) and 800℃ (lower than the melting point of Ti but much higher than the boiling point of Zn). The holding time for both was 30 minutes.

[0155] Results and Defect Analysis:

[0156] At 420℃: the Zn layer melts, but since the Ti layer remains solid, only a simple solid-liquid contact can be formed. Under the action of surface tension, the molten Zn cannot be stably maintained on the surface of the porous Ti framework, and it aggregates into spheres and flows, severely damaging the designed pore structure. The interface bonding area is a simple mechanical mosaic.

[0157] At 800℃, Zn not only melts completely, but its vapor pressure also rises sharply (boiling point is 907℃), causing it to evaporate and boil violently at high temperatures. This results in a large number of macroscopic pores and collapse in the porous Zn region, which separates from the Ti matrix. The interface strength is extremely low, and it crumbles with the slightest touch.

[0158] Comparative conclusion: Traditional heat treatment remelting cannot achieve localized heating and extremely fast thermal cycling. When dealing with materials with large melting and boiling point differences, it either fails to achieve effective metallurgical bonding or causes low-melting-point materials to evaporate or be lost due to excessive heat input, which completely fails to meet the requirements of precision porous components.

[0159] Experimental Example 1

[0160] The metal bone implants prepared in Examples 1-2 and Comparative Example 1 were subjected to performance testing experiments, including surface treatment, material characterization, material cell compatibility testing, and in vivo osteogenic activity verification.

[0161] 1. Surface treatment: The implant was ultrasonically cleaned in acetone, ethanol and deionized water for 10 min to remove loose powder remaining in the scaffold. After the sample was air-dried in an oven at 37℃, the scaffold was polished in a solution containing 5% hydrochloric acid, 5% nitric acid and 90% anhydrous ethanol for 2 min. Then the sample was ultrasonically cleaned in a solution of 10% nitric acid and 90% anhydrous ethanol for 10 s to remove residual Zn particles on the material surface. Finally, it was ultrasonically vibrated in anhydrous ethanol for 5 min to remove residual acid.

[0162] 2. Material Characterization

[0163] The surface and cross-sectional morphology and elemental distribution of the material were observed using a scanning electron microscope (SEM) equipped with an energy dispersive spectroscopy (EDS) instrument. Samples with dimensions of φ6.0 mm × 6.0 mm were prepared and subjected to compression tests at 25 °C using a universal testing machine with a displacement rate set to 0.1 mm / s.

[0164] Immersion tests were conducted according to ASTM-G31-72 using samples with a diameter of φ6.0mm × 6.0mm. The tests were performed for 28 days in a Hanks equilibrium salt solution at 37°C, maintaining a concentration of 20 mL / cm³. 2 The solution-to-surface area ratio was determined. The Hanks equilibrium salt solution was replaced every 3 days. After immersion, the surface and cross-sectional morphology of the samples were observed using a scanning electron microscope equipped with an energy dispersive spectroscopy (EDS) instrument, and corrosion products were identified using X-ray diffraction at a scan rate of 2° / min and a step size of 0.02°, with a scan range of 10° to 80°. The samples were then ultrasonically cleaned with a 10% chromic acid solution for 5 min to remove corrosion products. The corrosion rate (CR) was calculated using the following formula:

[0165]

[0166] Where C is the degradation rate, mm / a; K = 8.76 × 10⁻⁶ 4 ΔW is the mass loss, in g; A is the surface area of ​​the sample, in cm². 2 D is the sample density, g / cm³ 3 T represents the soaking time, in hours.

[0167] Finally, in order to determine the Zn released in the sample 2+ The concentration was determined by immersing the sample in HBSS for 3, 7, 14, and 28 days without changing the solution. At the specified time points, the Zn concentration in the HBSS was measured by inductively coupled plasma mass spectrometry. 2+ The concentration.

[0168] 3. Material cell compatibility testing

[0169] Mouse osteoblast precursor cells (MC3T3-E1) were cultured in α-MEM complete medium at 37°C, 5% CO2, and 95% humidification. This medium contained 10% fetal bovine serum and 1% penicillin-streptomycin solution. Cells were passaged when they reached 80%-90% confluence. Cells from passages 3-8 were used in this invention. Sample extracts were obtained according to ISO 10993-5. Samples with a diameter of φ6.0 mm × 6.0 mm were sterilized by Co-60 irradiation and then immersed in α-MEM complete medium at 37°C for 24 ± 0.5 h in a cell culture incubator, achieving an extraction rate of 1.25 cm⁻¹. 2 / mL. Zn in the sample extract was determined by ICP-MS. 2+ concentration.

[0170] Cell counting kit-8 was used to assess the viability of MC3T3-E1 cells. Cells were counted at 2 × 10⁸ cells / year. 3 Cells were seeded at a density of 1 / 2 wells into 96-well plates until the cells adhered to the bottom of the wells. The original culture medium was then replaced with sample extract, and the cells were cultured for 1, 3, and 5 days. At specific time points, 10 μL of LCK-8 solution was added, and the cells were incubated at 37°C in the dark for 1 hour. The optical density (OD) value was measured at 450 nm using a multi-sensor microplate reader. Finally, cell viability was calculated using the following formula: Cell viability (%) = [(Experimental group OD value - Blank group OD value) / (Control group OD value - Blank group OD value)] × 100%. For live / dead cell staining, MC3T3-E1 cells were co-cultured with 10% and 30% extracts. Cells were stained using a live / dead cell staining kit, and cell viability and morphology were assessed by fluorescence microscopy.

[0171] The adhesion morphology and distribution of MC3T3-E1 cells on the sample surface were observed using SEM. Sterile samples were co-cultured with MC3T3-E1 cells for 24 h. Then, the samples were washed with PBS, fixed with 4% paraformaldehyde fixative, and gradually dehydrated using a gradient of 30%, 50%, 70%, 80%, 90%, 95%, 100%, and 100% ethanol. Finally, the samples were dried in a vacuum drying oven, and the cells on the samples were observed.

[0172] 4. Verification of osteogenic activity of the material in vivo

[0173] Samples measuring φ2.8mm × 4.8mm were selected for in vivo implantation. Male rats weighing 200-250g were anesthetized by intraperitoneal injection of ketamine (10mg / kg) and 2% toluidine (10mg / kg). A longitudinal incision was made lateral to the patellar ligament to expose the lateral femoral condyle. A single cortical perforation bone defect model of the rat lateral femoral condyle was established by drilling a hole with a 2.4mm drill bit. The sample was implanted into the defect with the Zn portion facing inward. The surgical site was then sutured layer by layer. Rats were euthanized at 4 and 12 weeks post-implantation to collect the femur.

[0174] Rat femurs were collected and fixed with 4% paraformaldehyde solution. The fixed femurs were scanned using a high-resolution X-ray computed tomography scanner and imaged using the 3D image processing software CTvox 3.0. The region of interest (ROI) was defined as a cylindrical region of φ3.3 mm × 5.2 mm, with the sample located at the center of the ROI. Bone volume fraction (BV / TV), trabecular number (Tb.N), trabecular thickness (Tb.Th), and trabecular separation (Tb.Sp) in the ROI were quantitatively analyzed to assess bone formation.

[0175] Results Analysis

[0176] 1. After acid washing, the sample generally showed no excess powder in the Zn region (see [link to sample description]). Figure 2 The scanning electron microscope results show (see below) Figure 4 The Zn region surface was smooth with no unmelted powder adhering to it. No obvious cracks were observed in the interfacial bonding region of the Zn / Ti composite, indicating good bonding between Zn and Ti. Notably, a small amount of Ti was found in the Zn region near the interface, which may be due to the disturbance of the molten pool during Zn printing, causing liquid Ti to be sputtered into the Zn region.

[0177] 2. Please see Figure 5 The compressive strength (σbc) of pure Zn and 1-Zn / Ti is 210.38±10.15MPa and 423.10±25.35MPa, respectively. 3-Zn / Ti exhibits the highest σbc, which is 505.73±10.15MPa.

[0178] 3. Please see Figure 6 a. The in vitro corrosion rate of pure Zn was approximately 1.68 ± 0.12 mm / y, the corrosion rate of 1-Zn / Ti was 5.10 ± 0.57 mm / y, and the corrosion rate of 3-Zn / Ti accelerated to approximately 10.39 ± 0.87 mm / y, which was 6.1 times that of pure Zn. Furthermore, this invention also monitored the release of Zn from samples in HBSS. 2+ Concentration. Please see [link / reference]. Figure 6b, In the first 3 days, Zn in all samples 2+ The release is very rapid. The Zn released from pure Zn... 2+ The surface gradually stabilized after 3 days, likely due to the formation of corrosion products on the sample surface, which hindered the contact between the corrosive medium and the Zn matrix. However, despite the presence of more corrosion products on the Zn / Ti composite surface, the released Zn... 2+ The concentration continued to increase. Furthermore, Zn was released from the 3-Zn / Ti group. 2+ The concentrations were consistently higher than those of 1-Zn / Ti. These results indicate that the degradation rate of Zn / Ti composites is faster than that of pure Zn, with 3-Zn / Ti exhibiting the fastest degradation rate.

[0179] The amount of corrosion products is closely related to the degradation rate of Zn; the more corrosion products, the more severe the degradation behavior. XRD patterns of the degradation products of the samples after 28 days of immersion are shown in the attached image. Figure 6 c) The presence of Zn(OH)2, ZnO, and Zn3(PO4)2·4H2O was confirmed. Furthermore, the diffraction peak intensity of the corrosion products in the Zn region of the 3-Zn / Ti sample was significantly stronger than that of pure Zn and 1-Zn / Ti. In addition, the scanning electron microscopy results of the degraded sample surface (see [link to relevant documentation]) Figure 6 d) shows that more white corrosion products were observed in the Zn region of the Zn / Ti sample compared to pure Zn, with the highest accumulation of products on the Zn surface of the inner Zn layer in 3-Zn / Ti. EDS results show that the sample surface is significantly enriched with Zn, Ca, and P elements, which means that Zn oxides or hydroxides, calcium compounds, and phosphates are present in the corrosion products.

[0180] 4. Please see Figure 7 a. Zn in pure Zn extract 2+ The concentration was approximately 24.49 ± 3.63 μg / mL, and the Zn concentration in the 1-Zn / Ti extract was... 2+ The concentration increased to 70.46±8.86 μg / mL, and the Zn concentration in the 3-Zn / Ti extract increased. 2+ The highest concentration was 103.87 ± 3.31 μg / mL. Figure 7 b presents the CCK-8 results, showing that the pure Zn group exhibited good cell compatibility in all diluted extracts, promoting cell proliferation at 50% and 25% concentrations, while at 12.5% ​​and 6.25% concentrations, due to Zn... 2+ At concentrations too low, there was no significant ability to promote cell proliferation. The extracts of the 1-Zn / Ti and 3-Zn / Ti groups at concentrations of 50% and 25% showed no significant effect on promoting cell proliferation due to Zn... 2+Excessive concentrations exhibited significant cytotoxicity. At a concentration of 12.5%, it showed grade 0 cytotoxicity and promoted cell proliferation, meeting the requirements for cytotoxicity assessment using extracts of biodegradable implant materials diluted 6-10 times. At 5 days, cell viability in the two groups reached approximately 124% and 126%, respectively, indicating that Zn... 2+ At low concentrations, it exhibits good cell compatibility. However, at a concentration of 6.25%, the 3-Zn / Ti group shows higher cell compatibility than the 1-Zn / Ti group because the Zn2+ ion concentration is still within the threshold for promoting cell proliferation.

[0181] Live / dead cells were stained using 25% and 12.5% ​​concentration extracts, and the results are as follows: Figure 7 As shown in c. At a concentration of 25%, the number of cells in the Zn group increased. In contrast, the 1-Zn / Ti and 3-Zn / Ti groups showed significant cytotoxicity. Specifically, the 1-Zn / Ti group exhibited a reduced number of viable cells and poor cell morphology, while the 3-Zn / Ti group showed an increased proportion of dead cells, with cells shrinking into spherical shapes. Conversely, at a concentration of 12.5%, the extracts from the 1-Zn / Ti and 3-Zn / Ti groups significantly promoted cell proliferation, with cells exhibiting normal stretched morphology, including spindle-shaped, irregularly triangular, and visible pseudopodia. The 3-Zn / Ti group showed a greater number of cells and better morphology. Cell morphology observed by SEM ( Figure 7 d) This further confirmed the viability results. On the Zn region surfaces of Zn, 1-Zn / Ti, and 3-Zn / Ti, MC3T3-E1 cells showed some shrinkage, but pseudopodia stretching was also observed, indicating some contact inhibition. In contrast, the cell morphology on the Ti region scaffold surfaces of 1-Zn / Ti and 3-Zn / Ti was slightly improved, with cell spreading and pseudopodia stretching. This may be due to the Zn... 2+ The long-distance diffusion and stimulation by microcurrents accelerate cell spreading.

[0182] 5. For example Figure 8 As shown in Figure a, a small amount of new bone was observed around all samples at week 4 post-surgery. The amount of new bone gradually increased over time, and at week 12, significantly more new bone was observed around the 1-Zn / Ti and 3-Zn / Ti groups compared to pure Zn. Quantitative analysis of osteogenic-related indices further confirmed the enhanced osteogenic capacity of the Zn / Ti samples, as higher BV / TV, Tb.Th, and Tb.N were observed (see Figure a). Figure 8 b) The enhanced osteogenic capacity may be due to Zn 2+ The long-distance diffusion and the presence of potential difference in 1-Zn / Ti and 3-Zn / Ti.

[0183] The above results demonstrate that heterogeneous zinc / titanium metal bone implants based on multi-material additive manufacturing, through optimized interfacial bonding and layered configuration design, achieve a synergistic improvement in mechanical strength (compressive strength >500 MPa) and degradation rate (10.39 mm / y) compared to pure Zn, approximately 2.5 and 6 times respectively. Upon contact with bodily fluids, they release Zn... 2+ At appropriate concentrations, it significantly promoted the proliferation of MC3T3-E1 cells (>124%), and the resulting potential difference is expected to mimic endogenous electrical signals. Its ability to promote new bone regeneration was validated in animal models (new bone formation increased by >2-fold). This material system provides an innovative solution for biodegradable bone implants that combines excellent mechanical properties, controllable degradation, and bioactivity through a triple mechanism of "heterogeneous metal-electrical couple effect-ion release".

[0184] Experiment Example 2

[0185] Samples with a diameter of 10.0 mm × 2.0 mm were sterilized by Co-60 irradiation, and then immersed in α-MEM complete culture medium at 37°C for 24 ± 0.5 h in a cell culture incubator, with an extraction rate of 1.25 cm⁻¹. 2 / mL. Mg in the sample extract was determined by ICP-MS. 2+ and Zn 2+ concentration.

[0186] Results Analysis

[0187] Mg in pure Mg extract 2+ The concentration was 107.73 ± 7.77 μg / ml, and the pure Zn group contained almost no Mg. 2+ The Mg / Zn group contained approximately 231.63 ± 6.45 μg / ml of Mg. 2+ (Please see) Figure 9 a) The pure Mg group contains almost no Zn. 2+ Zn in pure Mg extract 2+ The concentration was 14.65 ± 0.32 μg / ml, and the Zn concentration in the Mg / Zn group was approximately 0.29 ± 0.03 μg / ml. 2+ (Please see) Figure 9 b).

[0188] The results showed that, unlike the accelerated Zn degradation effect of the Zn / Ti combination in Examples 1 and 2, the Mg / Zn combination in Example 3 exhibited a different effect on the regulation of degradation rate compared to the expected results. Specifically, the ion release rate in the Mg region reached 2.3 times that of pure Zn, while the ion release in the Zn region was inhibited. This phenomenon may lead to the risk of gas production and mechanical failure in the Mg region due to excessive degradation after implantation; at the same time, the expected osteogenic effect of the Zn region may not be fully realized due to the inhibited degradation.

[0189] Experimental Example 3

[0190] Because Ti-20Zn, Ti-30Zn and Ti-90Zn in Comparative Example 2 exhibited excellent mechanical properties and biomechanical compatibility, performance testing experiments were conducted on them, including material characterization, material cell compatibility testing, and in vivo osteogenic activity verification.

[0191] 1. Material Characterization

[0192] The samples were mechanically ground and polished. The microstructure and chemical composition were characterized at 20 kV using field emission scanning electron microscopy (SEM, Mira3 LMH, Tescan, Czech Republic) and energy-dispersive X-ray spectroscopy (EDS, X-max 20, Oxford Instruments, UK). X-ray diffraction (XRD, SmartLab SE, Rigaku, Japan) was performed in the range of 10–90° at a scan rate of 5° / min and a scan step of 0.02°.

[0193] Static immersion tests were conducted by immersing the disc samples in SBF solution at 37 °C for up to 8 weeks. Three samples with dimensions of Φ8 mm × 2 mm were prepared. The solution-to-sample area ratio was set to 20 mL / cm². 2 According to ASTM-G31-72, inductively coupled plasma atomic emission spectrometry (ICP-OES; Agilent 5110, Agilent, USA) was used. To investigate degradation products, surface morphology was observed using SEM, elemental composition was analyzed using EDS, chemical functional groups were characterized using Fourier transform infrared spectroscopy (FTIR, Nicolet iN10, Thermo Fisher, USA), and the phase was studied using micro-XRD. Degradation products were then removed using a solution containing 250 g / L ammonium acetate according to GB / T 16545–2015. An analytical balance (CP224C; Ohaus, USA) was used before and after the immersion test. Based on the weight loss results, the degradation rate of the composite material at different immersion stages was calculated using the formula: c = Δm / (ρ·A·t), where c is the degradation rate (mm / year), Δm is the weight loss, ρ and A are the density and surface area, and t is the immersion duration.

[0194] 2. Material cell compatibility testing

[0195] The mouse osteoblast precursor cell line (MC3T3-E1) was selected to determine the cell compatibility of the TiZn composite material. To prepare the extract, the block was immersed in α-MEM for 24 hours, with a medium volume to sample area ratio of 1.25 mL / cm². 2The extracts were diluted to 50%, 25%, and 12.5% ​​for subsequent in vitro experiments. Pure Ti and pure Zn were used as controls, and the culture medium was used as a negative control. Cell viability was assessed using a cell counting kit-8 (CCK-8, Beyotime, China) after 1 and 3 days of incubation. To further verify the effect of the compound extracts on cell morphology, live / dead cell staining and cytoskeleton staining were performed.

[0196] 3. Verification of osteogenic activity of the material in vivo

[0197] Twenty male SD rats weighing 200-250 g were used in the animal experiments. A bone defect with a diameter of 3 mm was created using a drill to implant the material. Six surgical groups were established: the Ti and Ti-90Zn groups were sacrificed at 8 and 12 weeks, and the Ti-20Zn and Ti-30Zn groups were sacrificed at 12 weeks, with at least 3 rats in each group at each time point.

[0198] Healing status was assessed using micro-computed tomography (Micro-CT, Skyscan 1276, Bruker, Germany). Three parameters—bone mineral density (BMD), bone volume fraction (BV / TV), and trabecular bone thickness (Tb.Th)—were quantitatively analyzed using three-dimensional (3D) reconstruction. Micro-CT analysis was performed only on rats fed for 12 weeks. Hard tissue sections were then prepared and stained to assess new bone formation. Rat femurs were embedded in methyl methacrylate and cut along the femoral shaft (perpendicular to the implant axis). Massen staining was then performed, and the sections were imaged using a high-resolution microscope (Eclipse Ci, Nikon, Japan). A thin layer of gold was coated onto the cross-sections for observation by SEM and EDS to investigate in vivo degradation and osseointegration.

[0199] Results Analysis

[0200] SEM and EDS results showed that all samples exhibited a core-shell structure, detected as uniformly distributed titanium particles encased in a complex shell. The interfaces in different regions of all TiZn composites were well-bonded and continuous, indicating good load transfer capacity during mechanical deformation (see [link to relevant documentation]). Figure 3 b).

[0201] Due to the inherent limitations of powder sintering, the Ti-Zn composite material prepared in Comparative Example 2 exhibits significant interdiffusion and mixing of metal elements, preventing it from forming a heterogeneous structure with clear material interfaces and three-dimensional interconnected pores as in Invention Examples 1 and 2.

[0202] Degradation morphology at different soaking periods, such as Figure 10 As shown. Ti-20Zn (see also...) Figure 10 a) and Ti-30Zn (see also) Figure 10 a) Micropitting corrosion did not appear on the sample until 4 weeks later. For Ti-90Zn, the white degradation deposits originated from zinc-rich areas (see [link to sample description]). Figure 10 a) forming tentacles (see also) Figure 10 a) and petal-shaped (see also) Figure 10 a) Clusters, eventually forming macroscopic sediments. EDS diagram shown (see [link]). Figure 10 (b) The petal-shaped aggregates contain Zn, Ca, P, O and C elements. Figure 10 The FTIR spectrum of c shows that the main chemical group is PO4. 3− and CO3 2− Combining Figure 10 The XRD pattern in sample d confirms that the products are mainly phosphates and carbonates, including CaZn2(PO4)2, CaZn(CO3)2, and ZnO. The Zn ions released from the SBF solution are as follows... Figure 10 As shown in e, influenced by degradation and precipitation formation, it first increases and then decreases, reaching a peak at 2-4 weeks. Furthermore, the Zn ion concentration in Ti-90Zn is significantly higher compared to the other two groups. The weight of Ti-90Zn and pure Zn is also significantly reduced, while the others are negligible, such as... Figure 10 As shown in f. Based on the weight loss method, the degradation rate of Ti-90Zn was initially calculated to be 93 μm / year, then decreased to 15–27 μm / year, while pure zinc was 123 μm / year in the first week. In contrast, Ti-20Zn and Ti-30Zn showed slight degradation, with maximum degradation rates of 2.1 and 7.6 μm / year, respectively. A comparison of the degradation rates of Ti-Zn samples with typical Mg and Zn alloys is shown in f. Figure 10 As shown in g, the degradation rates of different materials are as follows: in the initial stage of immersion, Ti-90Zn < pure Zn; after a long period of immersion, Ti-90Zn ≈ biocompatible zinc-based alloy.

[0203] The degradation products of the Ti-Zn composite material in Comparative Example 2 are similar to those of the Zn / Ti heterostructure composite materials in Examples 1 and 2, mainly consisting of zinc oxides and phosphates. However, the early Zn degradation rate of Ti-Zn in Comparative Example 2 is lower than that of pure Zn, while the degradation rate of the Zn region in the Zn / Ti heterostructure of Examples 1 is significantly higher than that of pure Zn. This difference is attributed to the core-shell structure formed by Ti-Zn in Comparative Example 2, whose slower degradation rate is difficult to match with the osteogenic cycle, contradicting the design goal of this invention to accelerate Zn degradation to promote repair.

[0204] Ti-90Zn and pure Zn extracts at concentrations of 100% and 50% significantly inhibited cell viability due to their high degradation rates during the initial soaking stage. In contrast, pure Ti, Ti-20Zn, Ti-30Zn, and other dilution groups showed acceptable cell compatibility with cell viability exceeding 90% (see [link to original text]). Figure 11 a). Further explanation of live / dead cell and cytoskeleton staining (see [link to documentation]). Figure 11 (b) Cells cultured with the diluted extract had a healthy shape and diffusion state, and the experimental and control groups showed similar levels of green fluorescence and fully diffused red stretched actin filaments.

[0205] Typically, the degradation products of materials in a biological environment are diluted by the continuous exchange of bodily fluids. After 6-10 times dilution, the ion concentration of in vitro extracts can be comparable to that in the actual in vivo environment. Furthermore, Zn... 2+ Its effect on osteoblasts is bidirectional—it promotes proliferation and differentiation only within an appropriate concentration range; excessively high concentrations produce cytotoxicity, while excessively low concentrations weaken the promoting effect. Therefore, it can be inferred that in Comparative Example 2, Ti-Zn promotes proliferation and differentiation only within an appropriate concentration range; excessively high concentrations produce cytotoxicity, while excessively low concentrations weaken the promoting effect. 2+ The release amount is relatively small, and its extract, after equivalent dilution, will have a weaker ability to promote cell proliferation than the Zn / Ti materials in Examples 1 and 2 under the same dilution conditions. 2+ Concentration environment.

[0206] Figure 12 Figure a shows the experimental procedure flowchart and photographs of rat femurs. Ti-Zn samples were implanted after modeling, and bone repair was assessed by micro-CT and histology at 8 and 12 weeks. Pure titanium was selected as a control. From the macroscopic images and scanned views of the rat femurs, the implants maintained their shape during the experiment, without dislocation or severe degeneration. Figure 12 3D reconstruction of the implants in group b showed that more bone was generated around the implant with increasing zinc content. Cross-sectional views were obtained at the top and bottom quarter of the implants, revealing that the bone tissue around the implants in the Ti-30Zn and Ti-90Zn groups was thicker than the other two groups, with almost no visible growth gaps (yellow arrows). In contrast, Ti-20Zn, due to its lower degradation capacity, exhibited similar bone repair effects to pure Ti, resulting in poor function of the bioactive components. In the marginal regions of the Ti-30Zn and Ti-90Zn implants (magnified image), signs of degradation products and new bone traces were observed, indicating good integration between the implant and surrounding bone tissue. Furthermore, in… Figure 12In step c, the trabecular bone and bone mineral density (BMD) of the implant surface were quantitatively assessed. BMD is used to obtain the mineral content and density in bone, BV / TV is a commonly used indicator for assessing cortical and cancellous bone mass, and Tb.Th is an important parameter for assessing trabecular bone strength. CT analysis revealed that the values ​​of Ti-30Zn and Ti-90Zn were significantly higher than the other two groups, while there was no significant difference between Ti and Ti-20Zn. These results indicate that Ti-Zn composite materials with relatively high Zn content have potential as bone implants.

[0207] Massen staining was performed on the hard tissue sections to observe bone regeneration around the implant. Figure 12 d). All groups exhibited the same trend: a layer of fibrous tissue (FT) was present around the implant ring, with mature new bone (NB) mineralized and deposited on the outer layer and connected to existing bone tissue (EB). At 8 weeks of implantation, degradation products (DP) appeared around the Ti-90Zn implant, with the circumferential NB layer thicker than in the pure Ti group. This DP layer was also widely observed in pure Zn and Zn alloy scaffolds. After 12 weeks, FT was found in the pure Ti and Ti-20Zn groups, with mature bone tissue encapsulated by FT. In the Ti-30Zn and Ti-90Zn groups, the distribution of FT decreased significantly with the increase of newly formed mature bone tissue. The DP around the Ti-90Zn implant was more pronounced than at 8 weeks, with FT extending into the DP region, corresponding to micro-CT analysis. No inflammation or adverse reactions were observed around the implant.

[0208] SEM and EDS mapping images were captured to analyze the bone-implant interface 12 weeks after implantation. Figure 13 Dark, C-rich interstitial spaces exhibiting organic characteristics were observed under electron beam. Therefore, the transition zone at the interface was identified as a FT layer composed of collagen, consistent with histological analysis. The FT layer thickness gradually increased and showed improved osteogenic capacity. Compared to Ti-30Zn, the Ti-90Zn implant was surrounded by a distinct DP layer composed of Zn, Ca, P, O, and C. This is similar to the previously reported in vivo degradation behavior of pure Zn and Zn-Mg scaffolds, with the DP primarily composed of zinc phosphate and calcium zinc phosphate salts. The close contact between the implant and bone tissue indicates good osseointegration capacity of the Ti-Zn composite.

[0209] Although the Ti-Zn material in Comparative Example 2 exhibited a similar trend of new bone regeneration as in Examples 1 and 2, it still had significant limitations:

[0210] ① In Comparative Example 2, there are fibrous tissue septa between the new bone and the metal matrix, while in Examples 1 and 2, the new bone directly integrates with the material at the degradation products in both the Ti and Zn regions. It is speculated that this difference is due to the microcurrent effect generated by the heterogeneous structure in the examples, which simulates the endogenous electrical signal and thus promotes bone integration.

[0211] ② The Ti-Zn material in Comparative Example 2 cannot achieve the clearly defined functional partitioning structure as in the examples, making it difficult to further optimize its performance.

[0212] Experiment Example 4

[0213] Material characterization was performed on the metal bone implant prepared in Comparative Example 3.

[0214] The morphology and elemental distribution of the material cross section were observed using a scanning electron microscope (SEM) equipped with an energy dispersive spectroscopy (EDS) instrument.

[0215] Results Analysis

[0216] Scanning electron microscopy (SEM) results show that the Zn / Ti interface region exhibits an uneven structure, pores, and granular morphology. This morphology may reflect the physical morphology during material bonding, and the pores or granular defects may affect the interface's compactness and mechanical properties. Simultaneously, a large number of Ti particles are sputtered in the Zn region near the interface, making the Zn-Ti boundary indistinct (see [link to SEM]). Figure 14 ).

[0217] It is evident that when performing multi-material additive manufacturing of dissimilar materials with large differences in melting and boiling points, interfacial remelting can greatly improve the interfacial bonding of the materials.

[0218] Although the principles of the present invention have been described in detail above with reference to preferred embodiments, those skilled in the art should understand that the above embodiments are merely illustrative explanations of the implementation of the present invention and are not intended to limit the scope of the present invention. The details in the embodiments do not constitute a limitation on the scope of the present invention. Any obvious changes, such as equivalent transformations or simple substitutions, based on the technical solutions of the present invention without departing from the spirit and scope of the present invention fall within the protection scope of the present invention.

Claims

1. A method for preparing a heterogeneous metal bone implant, characterized in that, Includes the following steps: S1: Based on the three-dimensional image data of the bone defect site, reconstruct the bone defect model, and define the spatial distribution and geometric configuration of two or more different metal materials with large melting and boiling point differences in the bone defect model to obtain the solid model of the bone implant. S2: Based on the physical model of the bone implant, pore structure parameters are independently designed for the two different metal material regions, and an L-PBF printing file is generated to obtain a porous model of the bone implant. S3: Under the protection of inert gas, based on the L-PBF printing file, multi-material laser powder bed melting technology is used to perform multi-material L-PBF integrated printing using two or more different metal material powders with large melting and boiling point differences to obtain a porous heterogeneous bone implant blank. Finally, the porous heterogeneous bone implant blank is post-processed to obtain a heterogeneous metal bone implant. Among them, the two or more different metal materials with large melting and boiling point differences are selected from metal systems with biocompatibility and different electrode potentials. By forming an electrochemical dipole in the body fluid environment, microcurrents are generated to simulate endogenous bioelectric signals to accelerate bone integration and regulate the degradation behavior of different materials.

2. The method for preparing heterogeneous metal bone implants according to claim 1, characterized in that, In step S1, the two different metallic materials are Ti-based metals and Zn-based metals; The Ti-based metal is selected from one or more of pure Ti, Ti6Al4V alloy, or Ti-0.3Mo-0.8Ni alloy, and the Zn-based metal is selected from one or more of pure Zn, Zn-Mg alloy, or Zn-Li alloy.

3. The method for preparing heterogeneous metal bone implants according to claim 1, characterized in that, In step S3, when performing the multi-material L-PBF integrated printing, after switching from one metal material powder to another, a layer of metal material powder to be printed is laid on the surface of the previous printed metal material, and the interface remelting process is implemented according to the laser parameters of the layer of metal material to be printed.

4. The method for preparing heterogeneous metal bone implants according to claim 3, characterized in that, When the metal powder is a Ti-based metal powder, the laser parameters of the laser powder bed are: power 120-300W, scanning speed 200-600mm / s, layer thickness 20-50μm; and / or, When the metal material is Zn-based metal powder, the laser parameters of the laser bed are: power 30-120W, scanning speed 100-800mm / s, and layer thickness 20-40μm.

5. The method for preparing heterogeneous metal bone implants according to claim 3, characterized in that, The metal powder is spherical or nearly spherical, with a particle size distribution D. 50 The particle size D of the different metal powders used in the same printing process is 10-60 μm. 50 The difference is no more than 30μm.

6. The method for preparing a heterogeneous metal bone implant according to claim 3, characterized in that, In step S2, the pore structure parameters include pore unit type, pore unit size, rod diameter, and unit structure porosity; The pore unit type includes one or more of the following: structures based on triple periodic minimal surfaces, lattice structures, and random biomimetic structures; The size of the pore unit is 1-5mm; The rod diameter is 0.4-4 mm; The porosity of the unit structure is 40-95%.

7. The method for preparing a heterogeneous metal bone implant according to claim 1, characterized in that, In step S3, the inert gas is selected from one or more of nitrogen, argon, or neon; and / or, The post-processing includes wire cutting to remove the support structure, cleaning, sandblasting, abrasive flow polishing, heat treatment, and surface functional treatment.

8. A heterogeneous metal bone implant obtained by the preparation method according to any one of claims 1-7.

9. The heterogeneous metal bone implant according to claim 8, characterized in that, The heterogeneous metal bone implant has a porous structure with a three-dimensional interconnected pore network, and the pore structure parameters in different metal material regions can be independently designed and controlled. In the heterogeneous metal bone implant, the two different metal materials are directly bonded at the interface without gradient composition transition; The heterogeneous metal bone implant has a layered configuration of at least two layers.

10. The use of the heterogeneous metal bone implant of claim 8 in the preparation of orthopedic and / or dental metal implants.

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