Preparation method for improving the stress corrosion resistance of magnesium alloys for load-bearing bone implants

Abstract

This invention discloses a method for preparing magnesium alloys with improved stress corrosion resistance for load-bearing bone implants, comprising the following steps: S1, heating a Mg-Gd-Zn-Zr as-cast alloy ingot prepared by semi-continuous casting to 300-450℃ under a pyrite protective atmosphere, holding for 4-12 hours, and then water-cooling to obtain a solution-treated alloy; S2, extruding the solution-treated alloy into plates with a thickness of 2-5 mm; S3, subjecting the upper and lower surfaces of the extruded alloy plates to ultrasonic surface rolling treatment; S4, aging the ultrasonically rolled plates with the following process parameters: temperature 80-120℃, holding for 1-4 hours, and then air-cooling. The magnesium alloy prepared by this invention exhibits uniform degradation behavior in simulated human body fluids and has high stress corrosion resistance, maintaining mechanical integrity and support during service life, thus meeting the application requirements of load-bearing fracture sites.

Description

Technical Field

[0001] This invention relates to a magnesium alloy, and more particularly to a method for preparing a magnesium alloy for improving the stress corrosion resistance of load-bearing bone implants, belonging to the field of biomedical magnesium alloy preparation technology. Background Technology

[0002] Currently, the implant materials used in clinical fracture treatment are mainly high-elasticity stainless steel and titanium alloys, which easily create stress shielding, leading to osteoporosis in non-stress areas. Furthermore, these materials are non-bioactive and non-degradable, offering extremely limited benefits for fracture healing. A second surgery is required after fracture healing to remove them, increasing the patient's surgical risks, psychological burden, and financial burden. Therefore, clinical expectations for internal fixation materials for bone trauma mainly focus on: good biocompatibility, with no hemolysis or coagulation reactions; good biosafety, with no carcinogenicity, teratogenicity, or sensitization; sufficient mechanical properties, with the material's elastic modulus close to that of bone tissue to reduce stress shielding; biodegradable and absorbable implants to avoid secondary surgery; and good bioactivity to promote osteogenesis.

[0003] Compared with traditional medical implant materials, alloys have many advantages in the treatment of fractures and bone defects, such as biodegradability, good biocompatibility and biomechanical compatibility, and density and elastic modulus close to human bone. They can not only effectively avoid the risks and burdens of "stress shielding" and secondary surgery for patients, but also promote osteogenic effects and accelerate fracture recovery, and have broad application prospects in the field of orthopedic fixation.

[0004] Although biodegradable pure magnesium and magnesium alloy bone fixation products are currently available, they all have certain shortcomings: pure magnesium, while having a slow degradation rate, has relatively low strength; magnesium alloys, due to the strengthening effect of alloying elements, have significantly improved strength, but the introduction of other phases and the large corrosion potential differences between different phases accelerate degradation and lead to uneven degradation. Furthermore, during service, magnesium alloys are prone to stress corrosion cracking due to the combined effects of external stress and corrosive media, causing them to prematurely lose structural integrity and mechanical support during degradation. Therefore, there is an urgent need to develop magnesium alloys for orthopedic implants with high resistance to stress corrosion (low stress sensitivity coefficient).

[0005] Surface modification can improve the corrosion resistance of magnesium alloys, but if the magnesium alloy matrix itself exhibits localized corrosion, surface treatment cannot change its localized corrosion characteristics (References: Fischerauer, SF; Kraus, T.; Wu, X.; Tangl, S.; Sorantin, E.). AC; JF; Uggowitzer, PJ; Weinberg, AM. In vivo degradation performance of micro-arc-oxidized magnesium implants: A micro-CT study in rats. Acta Biomater. 2013, 9, 5411-5420.

[0006] Degradation behavior of LAE442-based plate–screw-systems in an in vitro bonemodel.Mater.Sci.Eng.C.2015,49,305-315.

[0007] [https: / / doi.org / 10.1016 / j.msec.2015.01.019; Jiang, JH; Geng, X.; Zhang, XB](https: / / doi.org / 10.1016 / j.msec.2015.01.019; Jiang, JH; Geng, X.; Zhang, XB)](https: / / doi.org / 10.1016 / j.jma.2023.05.011). Therefore, pure magnesium and magnesium alloys can currently only be used clinically in non-load-bearing fracture sites, thus limiting their widespread clinical application.

[0008] Chinese invention patent application CN 112609068A discloses a composite strengthening method for improving the stress corrosion resistance of lightweight alloys. This composite strengthening method includes steps such as surface pretreatment of the lightweight alloy, pulsed magnetic field treatment, cryogenic laser shot peening, ultrasonic cleaning followed by ion implantation, and surface rolling strengthening treatment, thereby improving the stress corrosion resistance of the lightweight alloy. Although this invention involves surface rolling, the entire process is relatively complex and does not apply to the field of biomaterials and medical devices. Summary of the Invention

[0009] The purpose of this invention is to overcome the problem in the prior art that biomedical magnesium alloys suffer from stress corrosion cracking due to poor stress corrosion resistance, resulting in premature loss of structural integrity and mechanical support during service. This invention provides a method for preparing magnesium alloys with improved stress corrosion resistance for load-bearing bone implants. The magnesium alloy prepared by this method exhibits uniform degradation behavior in simulated human body fluids and possesses high stress corrosion resistance, maintaining mechanical integrity and support during service life, thus meeting the application requirements at load-bearing fracture sites.

[0010] To solve the above technical problems, the present invention provides a method for preparing magnesium alloy for improving stress corrosion resistance in load-bearing bone implants, comprising the following steps:

[0011] S1. The Mg-Gd-Zn-Zr as-cast alloy ingot prepared by semi-continuous casting is heated to 300-450℃ under a pyrite protective atmosphere and held for 4-12h, and then water-cooled to obtain a solution-treated Mg-Gd-Zn-Zr alloy.

[0012] S2. Extrude the solution-treated Mg-Gd-Zn-Zr alloy into plates with a thickness of 2-5 mm;

[0013] S3. The upper and lower surfaces of the extruded Mg-Gd-Zn-Zr alloy sheet are subjected to ultrasonic surface rolling treatment.

[0014] S4. The ultrasonically rolled sheet is subjected to aging treatment with the following process parameters: temperature 80-120℃, heat preservation for 1-4 hours, and then air cooling.

[0015] Preferably, in step S1, the mass percentage of each element in the Mg-Gd-Zn-Zr as-cast alloy ingot is: 1.5-6.4% Gd, 0.4-1.2% Zn, 0.2-0.8% Zr, with the remainder being Mg.

[0016] Preferably, in step S2: the temperature of the sheet and the die during extrusion is 280-380℃, the extrusion ratio is 8-25, and the extrusion speed is 10-80mm / s.

[0017] Preferably, in step S3, the process parameters for ultrasonic surface rolling treatment are: load 0.2-0.6 MPa, rolling amount 0.1-0.3 mm, linear speed 100-600 mm / min, step size 0.05-0.2 mm, ultrasonic frequency 20-50 kHz, and ultrasonic amplitude 5-20 μm.

[0018] Preferably, the magnesium alloy microstructure obtained in step S4 is divided into a fine-grained core region and ultra-fine-grained upper and lower surface regions, wherein:

[0019] The microstructure of the fine-grained core region consists of an α-Mg matrix, lamellar nanoscale LPSO and SFs distributed at the edge of the α-Mg matrix, and uniformly dispersed nanoscale precipitates.

[0020] The thickness of the upper and lower surface ultrafine crystalline regions is 100-300 μm, and they are composed of an α-Mg matrix, lamellar nanoscale LPSO and SFs distributed at the edge of the α-Mg matrix, uniformly dispersed nanoscale precipitates, and deformed twins in the matrix grains.

[0021] The nanoscale precipitates include Mg-Gd binary phase and Mg-Gd-Zn ternary phase.

[0022] Preferably, the magnesium alloy obtained in step S4 has a tensile strength of not less than 329 MPa in simulated human body fluid, a stress sensitivity coefficient of not more than 0.03, and an electrochemical corrosion current density of not more than 8.9 × 10⁻⁶. -6 A / cm 2 And not less than 4.4×10 -6 A / cm 2 .

[0023] Preferably, in step S1, the heating temperature of the ingot is 400℃ and the holding time is 6h.

[0024] Preferably, in step S2: the extrusion temperature is 320℃, the extrusion ratio is 16, and the extrusion speed is 60mm / s.

[0025] Preferably, in step S3: the ultrasonic surface rolling treatment has a load of 0.4 MPa, a rolling amount of 0.2 mm, a linear speed of 300 mm / min, a step size of 0.1 mm, a frequency of 30 kHz, and an amplitude of 10 μm.

[0026] Preferably, in step S4, the aging treatment temperature is 90–110°C, and the holding time is 1.5–3 hours.

[0027] Compared with existing technologies, this invention achieves the following beneficial effects: 1. The alloy's microstructure is divided into a core fine-grained region and upper and lower surface ultra-fine-grained regions; wherein, the core fine-grained region: the microstructure consists of a fine-grained α-Mg matrix, lamellar nanoscale long-period stacked ordered structures (LPSO) and stacking faults (SFs) distributed at the edges of the α-Mg matrix, and uniformly dispersed nanoscale precipitates. LPSO and SFs hinder dislocation movement and improve material strength; the uniformly dispersed nanoscale precipitates provide precipitation strengthening; and the fine-grained matrix microstructure provides fine-grain strengthening. Furthermore, the lamellar nanoscale thickness of LPSO and SFs distributed at the edge of the α-Mg matrix, along with the uniformly dispersed nanoscale precipitates, can effectively improve the corrosion resistance of the alloy. The mechanism is as follows: LPSO and SFs are rich in Gd and Zn atoms, while the precipitates, being Mg-Gd binary and Mg-Gd-Zn ternary phases, are also rich in Gd or Gd and Zn atoms, exhibiting superior corrosion resistance compared to the matrix. Their corrosion products are mainly Gd-containing oxides and hydroxides. When LPSO and SFs are distributed at the matrix edge, they form a relatively dense corrosion product layer, effectively hindering matrix corrosion and achieving uniform corrosion. When the nanoscale precipitates are dispersed, their corrosion products are also Gd-containing oxides and hydroxides, forming a relatively dense corrosion product layer, thus providing some protection to the matrix and improving the alloy's corrosion resistance and achieving uniform corrosion. Uniform corrosion can reduce stress concentration caused by localized pitting corrosion, thereby significantly improving stress corrosion resistance.

[0028] The upper and lower surface ultrafine-grained regions have a finer microstructure than the core region. In addition to the lamellar nanoscale thickness of LPSO and SFs distributed at the edge of the α-Mg matrix and the uniformly dispersed nanoscale precipitates, twinning structures are also formed. Due to the strengthening of fine grains and twinning, the strength of the alloy is further improved. The ultrafine-grained structure can further improve the corrosion resistance and stress corrosion resistance of the alloy.

[0029] 2. Layered structure offers the following advantages: For biodegradable materials used in load-bearing bone implants, a lower corrosion rate is not necessarily better. The supporting force (strength) provided by the implant material gradually decreases during degradation; conversely, the force (strength) it can withstand gradually increases with bone growth and healing. Ideally, the strengths of these two aspects are complementary, achieving a mechanical fit. Once the bone heals, the implant material is completely degraded. Therefore, clinically, it is desirable for the implant material to degrade slowly in the early stages of implantation to provide sufficient supporting force, while degradation accelerates in the later stages. In terms of degradation (corrosion) performance, the surface ultrafine-grained structure ensures that the alloy degrades slowly in the early stages of implantation; due to the uniform corrosion characteristics of the alloy, once the surface ultrafine-grained region is degraded, the core degrades relatively quickly, achieving a mechanical fit with the newly formed bone.

[0030] 3. Optimal solution treatment and extrusion processes can obtain a fine, uniform, fully recrystallized microstructure. This microstructure contains long-period stacked ordered phases (LPSO) rich in Gd and Zn atoms, with lamellar nanoscale thickness, and stacking faults (SFs) structures. These LPSOs and SFs are distributed at the edges of the α-Mg matrix, along with uniformly dispersed nanoscale precipitates, ensuring uniform corrosion of the alloy. The formation of LPSOs and SFs is closely related to stacking fault energy. In addition to alloy composition, the preparation and processing technology must be strictly controlled, including solution temperature, time, extrusion temperature, extrusion ratio, and extrusion rate. Improper process control may result in the absence of LPSOs and SFs, or their formation penetrating the matrix grains instead of being distributed at the grain edges. When the process control is appropriate, the alloy possesses a certain stacking fault energy, and Gd and Zn atoms aggregate at the matrix edges to form SFs. When both compositional and structural order are satisfied, some SFs transform into LPSOs. LPSO and SFs are rich in Gd and Zn atoms, and their corrosion resistance is better than that of the matrix. Their corrosion products are mainly Gd-containing oxides and hydroxides. When LPSO and SFs are distributed at the edge of the matrix, they can form a relatively dense corrosion product layer, which can effectively hinder the corrosion of the matrix and achieve uniform corrosion, thereby improving the stress corrosion resistance of the alloy.

[0031] 4. Ultrasonic Surface Rolling: This method combines ultrasonic impact energy with static rolling to treat the surface of extruded Mg-Gd-Zn-Zr alloy sheets. A certain amplitude of ultrasonic frequency mechanical vibration is applied along the normal direction of the workpiece surface by a processing head. Under certain feed conditions, the working head transmits static pressure and ultrasonic impact vibration to the surface of the alloy sheet, generating a pressing effect that creates a certain thickness of ultrafine grain layer and deformed twins on both the upper and lower surfaces of the Mg-Gd-Zn-Zr alloy sheet, generating a certain compressive stress on the sheet surface. First, the ultrafine grains and twins, through grain refinement and twinning strengthening, both contribute to improving the sheet's strength. Second, the ultrafine grains further improve the sheet's corrosion resistance and uniform corrosion behavior. Third, as a load-bearing bone implant material, it is mainly subjected to tensile stress during service. The compressive stress generated on the sheet surface by this invention can offset some of the tensile stress, thereby significantly improving its resistance to stress corrosion. It should be emphasized that when the surface compressive stress is too small, the effect on improving the stress corrosion resistance of the alloy is not obvious; when the surface compressive stress is too large, it will aggravate the stress corrosion of the alloy.

[0032] 5. Relatively low-temperature aging treatment can, on the one hand, appropriately regulate the surface compressive stress of the alloy, maintaining it between 45 and 60 MPa. During service, this can offset some tensile stress and improve stress corrosion resistance. On the other hand, it can accelerate the dispersion precipitation of nano-sized precipitates in the alloy without causing significant grain growth, thereby improving the alloy's strength, corrosion resistance, and stress corrosion resistance. The nano-precipitates are Mg-Gd binary and Mg-Gd-Zn ternary phases, rich in Gd or Gd and Zn atoms. Their corrosion products are also Gd-containing oxides and hydroxides, which can form a relatively dense corrosion product layer, thus providing a certain degree of protection to the matrix and improving the alloy's corrosion resistance and uniform corrosion, further enhancing its stress corrosion resistance. Furthermore, due to precipitation strengthening, the mechanical properties of the alloy are also further improved after aging treatment. Therefore, relatively low-temperature aging treatment ultimately improves the alloy's mechanical properties, corrosion resistance, and stress corrosion resistance.

[0033] 6. The preparation method of the present invention is simple and effective. The tensile strength of the obtained magnesium alloy structure in simulated human body fluid is not less than 329 MPa, which can meet the strength requirements of the human body's load-bearing fracture site. The surface compressive stress is between 45 and 60 MPa, the stress sensitivity coefficient does not exceed 0.03, and the alloy is uniformly corroded, so that the alloy will not lose its structural integrity and mechanical support prematurely due to stress corrosion or uneven corrosion, ensuring good mechanical support effect during service.

[0034] The electrochemical corrosion current density does not exceed 8.9 × 10⁻⁶. -6 A / cm 2 And not less than 4.4×10 -6 A / cm 2 This allows the corrosion rate of the alloy to be controlled between 0.1 and 0.2 mm / y, which can meet the degradation rate requirements of the product in the human body, while avoiding excessive corrosion that would lead to insufficient load-bearing capacity. This is of great significance for expanding the clinical application of biomedical magnesium alloys in load-bearing fracture sites. Attached Figure Description

[0035] The present invention will be further described in detail below with reference to the accompanying drawings and specific embodiments. The drawings are provided for reference and illustration only and are not intended to limit the present invention.

[0036] Figure 1 This is an electron backscattering scan image of the magnesium alloy obtained in Example 1 of the present invention;

[0037] Figure 2 This is a transmission electron micrograph of the magnesium alloy obtained in Example 1 of the present invention. Detailed Implementation

[0038] To make the technical means, creative features, objectives and effects of this invention easier to understand, the invention will be further described below with reference to specific illustrations.

[0039] Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this invention pertains. The terminology used herein is for the purpose of describing particular embodiments only and is not intended to limit the scope of the invention.

[0040] Example 1

[0041] The present invention provides a method for preparing magnesium alloys for improving stress corrosion resistance in load-bearing bone implants, comprising the following steps:

[0042] S1. The Mg-3.2Gd-0.8Zn-0.4Zr as-cast alloy ingot prepared by semi-continuous casting is heated to 400℃ under a pyrite protective atmosphere, held for 6h, and then water-cooled to obtain the solution-treated alloy.

[0043] S2. The solution-treated Mg-3.2Gd-0.8Zn-0.4Zr alloy is extruded into a sheet with a thickness of 3.5mm. During extrusion, the temperature of the sheet and the die is 320℃, the extrusion ratio is 16, and the extrusion speed is 60mm / s.

[0044] S3. The upper and lower surfaces of the extruded Mg-3.2Gd-0.8Zn-0.4Zr alloy sheet are subjected to ultrasonic surface rolling treatment. The corresponding process parameters are: load 0.3MPa, rolling amount 0.15mm, linear speed 300mm / min, step size 0.1mm, frequency 30kHz, and amplitude 10μm.

[0045] S4. The ultrasonically rolled sheet is subjected to aging treatment with the following process parameters: temperature 110℃, heat preservation for 3 hours, and then air cooling to room temperature.

[0046] The electron backscattering scan image of the magnesium alloy casting in Example 1 is shown below. Figure 1 As shown, the microstructure of the alloy is divided into a surface ultrafine-grained region and a core fine-grained region. The surface ultrafine-grained region is approximately 180 μm thick and consists of an α-Mg matrix, lamellar nanoscale LPSO and SFs (such as...) distributed at the edges of the α-Mg matrix. Figure 2 As shown in the figure, the microstructure consists of uniformly dispersed nanoscale precipitates and deformed twins in the matrix grains, with grain sizes significantly smaller than those in the core. The microstructure of the fine-grained core region consists of an α-Mg matrix, lamellar nanoscale thicknesses of LPSO and SFs distributed at the edges of the α-Mg matrix, and uniformly dispersed nanoscale precipitates.

[0047] The magnesium alloy product of Example 1 was subjected to low strain rate (3.6 × 10⁻⁶) tests in air and simulated human body fluids. - 6 s -1 The stress corrosion test yielded tensile strengths of the magnesium alloy product in air and in simulated human body fluids of 336 MPa and 329 MPa, respectively, with a stress sensitivity coefficient of 0.02. Electrochemical testing of the magnesium alloy from Example 1 in simulated human body fluids showed a corrosion current density of 8.9 × 10⁻⁶. -6 A / cm 2 The surface compressive stress of the magnesium alloy in Example 1 was tested and found to be 45.2 MPa.

[0048] Comparative Example 1

[0049] Steps S3 and S4 are omitted, and the remaining steps are the same as in Example 1, that is, steps S1 and S2 in Example 1 are used to obtain the extruded Mg-3.2Gd-0.8Zn-0.4Zr alloy.

[0050] The magnesium alloy sample of Comparative Example 1 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluids. - 6 s -1 Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 249 MPa in air and 201 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.19. Electrochemical testing in simulated human body fluid showed a corrosion current density of 3.9 × 10⁻⁶. -5 A / cm 2 .

[0051] Comparative Example 2

[0052] Step S4 is omitted, and the remaining steps are the same as in Example 1, that is, steps S1, S2 and S3 in the example are used to obtain the ultrasonically surface-rolled Mg-3.2Gd-0.8Zn-0.4Zr alloy.

[0053] The magnesium alloy sample of Comparative Example 2 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluids. - 6 s -1 Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 314 MPa in air and 283 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.10. Electrochemical testing in simulated human body fluid showed a corrosion current density of 1.2 × 10⁻⁶. -5 A / cm 2 The surface compressive stress was tested and found to be 90.8 MPa.

[0054] Comparative Example 3

[0055] Except for step S3, all other steps are the same as in Example 1, i.e., steps S1, S2, and S4 in Example 1 are used. In step S3, the upper and lower surfaces of the extruded Mg-3.2Gd-0.8Zn-0.4Zr alloy sheet are subjected to ultrasonic surface rolling treatment. The corresponding process parameters are: load 0.1MPa, rolling amount 0.1mm, linear speed 100mm / min, step size 0.05mm, ultrasonic frequency 10kHz, and ultrasonic amplitude 5μm.

[0056] The magnesium alloy sample of Comparative Example 3 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluids. - 6 s -1 Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 286 MPa in air and 242 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.15. Electrochemical testing in simulated human body fluid showed a corrosion current density of 2.0 × 10⁻⁶ MPa. -5 A / cm 2 The surface compressive stress was tested and found to be 14.2 MPa.

[0057] Comparative Example 4

[0058] Except for step S3, all other steps are the same as in Example 1, i.e., steps S1, S2, and S4 in Example 1 are used. In step S3, the upper and lower surfaces of the extruded Mg-3.2Gd-0.8Zn-0.4Zr alloy sheet are subjected to ultrasonic surface rolling treatment. The corresponding process parameters are: load 0.8MPa, rolling amount 0.4mm, linear speed 100mm / min, step size 0.05mm, ultrasonic frequency 50kHz, and ultrasonic amplitude 25μm.

[0059] The magnesium alloy sample of Comparative Example 4 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluid. - 6 s -1 Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 372 MPa in air and 297 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.20. Electrochemical testing in simulated human body fluid showed a corrosion current density of 5.3 × 10⁻⁶. -5 A / cm 2 The surface compressive stress was tested and found to be 203 MPa.

[0060] Comparative Example 5

[0061] Except for step S4, all other steps are the same as in Example 1, that is, steps S1, S2 and S3 in Example 1 are used. In step S4, the board material that has undergone ultrasonic surface rolling treatment is subjected to aging treatment, and the process parameters are: temperature 50°C, heat preservation for 4 hours, and then air cooling to room temperature.

[0062] The magnesium alloy sample of Comparative Example 5 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluid. - 6 s -1 Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 321 MPa in air and 292 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.09. Electrochemical testing in simulated human body fluid showed a corrosion current density of 1.1 × 10⁻⁶. -5 A / cm 2 The surface compressive stress was tested and found to be 86.3 MPa.

[0063] Comparative Example 6

[0064] Except for step S4, all other steps are the same as in Example 1, that is, steps S1, S2 and S3 in Example 1 are used. In step S4, the board material that has undergone ultrasonic surface rolling is subjected to aging treatment, and the process parameters are: temperature 150℃, heat preservation for 1 hour, and then air cooling to room temperature.

[0065] The magnesium alloy sample of Comparative Example 6 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluids. - 6 s -1 Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 345 MPa in air and 296 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.14. Electrochemical testing in simulated human body fluid showed a corrosion current density of 1.0 × 10⁻⁶. -5 A / cm 2 The surface compressive stress was tested and found to be 8.7 MPa.

[0066] Example 2

[0067] The present invention provides a method for preparing magnesium alloys for improving stress corrosion resistance in load-bearing bone implants, comprising the following steps:

[0068] S1. First, the Mg-3.2Gd-0.8Zn-0.4Zr as-cast alloy ingot prepared by semi-continuous casting is heated to 400℃ under a pyrite protective atmosphere and held for 6 hours, and then water-cooled to obtain the solution-treated alloy.

[0069] S2. The solution-treated Mg-3.2Gd-0.8Zn-0.4Zr alloy is extruded into a sheet with a thickness of 3.5mm. During extrusion, the temperature of the sheet and the die is 320℃, the extrusion ratio is 16, and the extrusion speed is 60mm / s.

[0070] S3. The upper and lower surfaces of the extruded Mg-3.2Gd-0.8Zn-0.4Zr alloy sheet are subjected to ultrasonic surface rolling treatment. The corresponding process parameters are: load 0.4MPa, rolling amount 0.2mm, linear speed 300mm / min, step size 0.1mm, frequency 30kHz, and amplitude 10μm.

[0071] S4. The ultrasonically rolled plate is subjected to aging treatment with the following process parameters: temperature 100℃, heat preservation for 2 hours, and then air cooling.

[0072] The microstructure of the alloy in Example 2 is divided into a fine-grained core region and ultrafine-grained upper and lower surface regions. The fine-grained core region consists of an α-Mg matrix, lamellar nanoscale LPSO and SFs distributed at the edges of the α-Mg matrix, and uniformly dispersed nanoscale precipitates. The ultrafine-grained upper and lower surface regions are approximately 220 μm thick and consist of an α-Mg matrix, lamellar nanoscale LPSO and SFs distributed at the edges of the α-Mg matrix, uniformly dispersed nanoscale precipitates, and deformed twins within the matrix grains.

[0073] The magnesium alloy product of Example 2 was subjected to low strain rate (3.6 × 10⁻⁶) tests in air and simulated human body fluids. - 6 s -1 The stress corrosion test of the magnesium alloy product yielded tensile strengths of 361 MPa in air and 354 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.02. Electrochemical testing of the magnesium alloy in simulated human body fluid showed a corrosion current density of 4.4 × 10⁻⁶ MPa. -6 A / cm 2 The surface compressive stress of the magnesium alloy in Example 2 was tested and found to be 56.8 MPa.

[0074] Comparative Example 7

[0075] Except for step S3, all other steps are the same as in Example 2, i.e., steps S1, S2, and S4 in Example 2 are used. In step S3, the upper and lower surfaces of the extruded Mg-3.2Gd-0.8Zn-0.4Zr alloy sheet are subjected to ultrasonic surface rolling treatment. The corresponding process parameters are: load of 0.2MPa, rolling amount of 0.1mm, linear speed of 100mm / min, step size of 0.05mm, frequency of 20kHz, and amplitude of 5μm.

[0076] The magnesium alloy sample of Comparative Example 7 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluid. - 6 s -1 Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 303 MPa in air and 278 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.08. Electrochemical testing in simulated human body fluid showed a corrosion current density of 3.4 × 10⁻⁶. -5 A / cm 2 The surface compressive stress was tested and found to be 17.3 MPa.

[0077] Comparative Example 8

[0078] Except for step S4, all other steps are the same as in Example 2, that is, steps S1, S2 and 3 in Example 2 are used. In step S4, the board material that has undergone ultrasonic surface rolling is subjected to aging treatment, and the process parameters are: temperature 80℃, heat preservation for 1 hour, and then air cooling to room temperature.

[0079] The magnesium alloy sample of Comparative Example 8 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluids. - 6 s -1 Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 324 MPa in air and 305 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.06. Electrochemical testing in simulated human body fluid showed a corrosion current density of 2.7 × 10⁻⁶. -5 A / cm 2 The surface compressive stress was tested and found to be 88.4 MPa.

[0080] Example 3

[0081] The present invention provides a method for preparing magnesium alloys for improving stress corrosion resistance in load-bearing bone implants, comprising the following steps:

[0082] S1. First, the Mg-3.2Gd-0.8Zn-0.4Zr as-cast alloy ingot prepared by semi-continuous casting is heated to 400℃ under a pyrite protective atmosphere and held for 6 hours, and then water-cooled to obtain the solution-treated alloy.

[0083] S2. The solution-treated Mg-3.2Gd-0.8Zn-0.4Zr alloy is extruded into a sheet with a thickness of 3.5mm. During extrusion, the temperature of the sheet and the die is 320℃, the extrusion ratio is 16, and the extrusion speed is 60mm / s.

[0084] S3. The upper and lower surfaces of the extruded Mg-3.2Gd-0.8Zn-0.4Zr alloy sheet are subjected to ultrasonic surface rolling treatment. The corresponding process parameters are: load 0.5MPa, rolling amount 0.25mm, linear speed 300mm / min, step size 0.1mm, frequency 30kHz, and amplitude 10μm.

[0085] S4. The ultrasonically rolled sheet is subjected to aging treatment with the following process parameters: temperature 90℃, heat preservation for 1.5h, and then air cooling.

[0086] The microstructure of the alloy in Example 3 is divided into a fine-grained core region and ultrafine-grained upper and lower surface regions. The fine-grained core region consists of an α-Mg matrix, lamellar nanoscale LPSO and SFs distributed at the edges of the α-Mg matrix, and uniformly dispersed nanoscale precipitates. The ultrafine-grained upper and lower surface regions are approximately 240 μm thick and consist of an α-Mg matrix, lamellar nanoscale LPSO and SFs distributed at the edges of the α-Mg matrix, uniformly dispersed nanoscale precipitates, and deformed twins within the matrix grains.

[0087] The magnesium alloy product of Example 3 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluids. - 6 s -1 The stress corrosion test yielded tensile strengths of the magnesium alloy product in air and in simulated human body fluids of 385 MPa and 372 MPa, respectively, with a stress sensitivity coefficient of 0.03. Electrochemical testing of the magnesium alloy in simulated human body fluids showed a corrosion current density of 5.3 × 10⁻⁶. -6 A / cm 2 The surface compressive stress of the magnesium alloy in Example 3 was tested and found to be 59.8 MPa.

[0088] Comparative Example 9

[0089] Except for step S3, all other steps are the same as in Example 3, i.e., steps S1, S2, and 4 in Example 3 are used. In step S3, the upper and lower surfaces of the extruded Mg-3.2Gd-0.8Zn-0.4Zr alloy sheet are subjected to ultrasonic surface rolling treatment. The corresponding process parameters are: load of 0.6MPa, rolling amount of 0.6mm, linear speed of 600mm / min, step size of 0.2mm, frequency of 50kHz, and amplitude of 20μm.

[0090] The magnesium alloy sample of Comparative Example 9 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluids. - 6 s -1Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 396 MPa in air and 323 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.18. Electrochemical testing in simulated human body fluid showed a corrosion current density of 2.5 × 10⁻⁶ MPa. -5 A / cm 2 The surface compressive stress was tested and found to be 166 MPa.

[0091] Comparative Example 10

[0092] Except for step S4, all other steps are the same as in Example 3, that is, steps S1, S2 and 3 in Example 3 are used. In step S4, the board material that has undergone ultrasonic surface rolling treatment is subjected to aging treatment, and the process parameters are: temperature 130℃, heat preservation for 4 hours, and then air cooling to room temperature.

[0093] The magnesium alloy sample of Comparative Example 10 was subjected to low strain rates (3.6 × 10⁻⁶) in air and simulated human body fluids. - 6 s -1 Stress corrosion tests were conducted on the magnesium alloy sample, yielding tensile strengths of 392 MPa in air and 308 MPa in simulated human body fluid, with a stress sensitivity coefficient of 0.21. Electrochemical testing in simulated human body fluid showed a corrosion current density of 9.6 × 10⁻⁶. -6 A / cm 2 The surface compressive stress was tested and found to be 26.3 MPa.

[0094] The main performance characteristics of Examples 1 to 3 and Comparative Examples 1 to 10 are shown in the table below. UTS Air This represents the alloy's tensile strength in air, UTS SBF This represents the tensile strength of the alloy in simulated human body fluids, measured using an electronic tensile testing machine at 3.6 × 10⁻⁶. -6 s -1 Measured at low strain rates; I UTS This represents the stress sensitivity coefficient of the alloy, which is calculated using the following formula. The surface compressive stress of the alloy was measured using an X-ray diffractometer; the corrosion current density was obtained by fitting the polarization curves after testing with an electrochemical workstation.

[0095]

[0096] The above description is merely a preferred embodiment of the present invention, showing and describing the basic principles, main features, and advantages of the present invention. It is not intended to limit the scope of patent protection of the present invention. Those skilled in the art should understand that the present invention is not limited to the above embodiments. In addition to the above embodiments, the present invention may have other implementations without departing from the spirit and scope of the invention. Various changes and modifications to the present invention are possible, and all technical solutions formed by equivalent substitutions or equivalent transformations fall within the scope of protection claimed by the present invention. The scope of protection of the present invention is defined by the appended claims and their equivalents. Technical features not described in the present invention can be implemented by or using existing technology, and will not be elaborated here.

Claims

1. A method for preparing magnesium alloys for improving stress corrosion resistance in load-bearing bone implants, characterized in that, The steps are as follows: S1. The Mg-Gd-Zn-Zr as-cast alloy ingot prepared by semi-continuous casting is heated to 300-450℃ under a pyrite protective atmosphere and held for 4-12h, and then water-cooled to obtain a solution-treated Mg-Gd-Zn-Zr alloy. S2. Extrude the solution-treated Mg-Gd-Zn-Zr alloy into plates with a thickness of 2-5 mm; S3. The upper and lower surfaces of the extruded Mg-Gd-Zn-Zr alloy sheet are subjected to ultrasonic surface rolling treatment. S4. The ultrasonically rolled sheet is subjected to aging treatment with the following process parameters: temperature 90~110℃, heat preservation for 1.5~3h, followed by air cooling. In step S1, the mass percentage of each element in the Mg-Gd-Zn-Zr as-cast alloy ingot is: 1.5-6.4% Gd, 0.4-1.2% Zn, 0.2-0.8% Zr, with the remainder being Mg; In step S3, the process parameters for ultrasonic surface rolling treatment are: load 0.3-0.6 MPa, rolling amount 0.1-0.3 mm, linear speed 100-600 mm / min, step size 0.05-0.2 mm, ultrasonic frequency 20-50 kHz, and ultrasonic amplitude 5-20 μm.

2. The preparation method for improving the stress corrosion resistance of magnesium alloy for load-bearing bone implants according to claim 1, characterized in that, In step S2: the temperature of the sheet and the die during extrusion is 280-380℃, the extrusion ratio is 8-25, and the extrusion speed is 10-80mm / s.

3. The preparation method for improving the stress corrosion resistance of magnesium alloy for load-bearing bone implants according to claim 2, characterized in that: The magnesium alloy microstructure obtained in step S4 is divided into a fine-grained core region and ultra-fine-grained upper and lower surface regions, wherein: The microstructure of the fine-grained core region consists of an α-Mg matrix, lamellar nanoscale LPSO and SFs distributed at the edge of the α-Mg matrix, and uniformly dispersed nanoscale precipitates. The thickness of the upper and lower surface ultrafine crystalline regions is 100-300 μm, and they are composed of an α-Mg matrix, lamellar nanoscale LPSO and SFs distributed at the edge of the α-Mg matrix, uniformly dispersed nanoscale precipitates, and deformed twins in the matrix grains. The nanoscale precipitates include Mg-Gd binary phase and Mg-Gd-Zn ternary phase.

4. The preparation method for improving the stress corrosion resistance of magnesium alloy for load-bearing bone implants according to claim 1, characterized in that: The magnesium alloy obtained in step S4 has a tensile strength of not less than 329 MPa in simulated human body fluid, a stress sensitivity coefficient of not more than 0.03, and an electrochemical corrosion current density of not more than 8.9 × 10⁻⁶. -6 A / cm² and not less than 4.4×10 -6 A / cm².

5. The preparation method for improving the stress corrosion resistance of magnesium alloy for load-bearing bone implants according to claim 1, characterized in that: In step S1, the heating temperature of the ingot is 400℃ and the holding time is 6h.

6. The preparation method for improving the stress corrosion resistance of magnesium alloy for load-bearing bone implants according to claim 1, characterized in that, In step S2: the extrusion temperature is 320℃, the extrusion ratio is 16, and the extrusion speed is 60mm / s.

7. The preparation method for improving the stress corrosion resistance of magnesium alloy for load-bearing bone implants according to claim 2, characterized in that: In step S3: the ultrasonic surface rolling treatment has a load of 0.4 MPa, a rolling amount of 0.2 mm, a linear speed of 300 mm / min, a step size of 0.1 mm, a frequency of 30 kHz, and an amplitude of 10 μm.

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