Degradable zinc alloy wire for small ear malformation correction and preparation method of degradable zinc alloy wire

Abstract

The invention discloses a degradable zinc alloy wire for small ear malformation correction and a preparation method of the degradable zinc alloy wire, and belongs to the field of biomedical materials. A degradable zinc alloy wire material for microtympanic malformation correction is characterized in that the degradable zinc alloy wire material is Zn-Cu-Mn-Fe zinc alloy, and the alloy elements of the degradable zinc alloy wire material comprise, by mass, 0 < Cu < = 3 wt.%, 0 < Mn < = 1 wt.%, 0 < Fe < = 1 wt.%, and the balance Zn. The alloy wire is prepared through the working procedures of smelting, casting, backward extrusion, cold drawing combined with hot drawing, heat treatment and the like. According to the Zn-Cu-Mn-Fe quaternary alloy wire obtained through the method, the tensile strength ranges from 100 MPa to 150 MPa, the yield strength ranges from 80 MPa to 120 MPa, the ductility is larger than 58%, the excellent plasticity, the appropriate tensile strength and the appropriate yield strength needed by a wire for microauricular malformation correction suture are achieved, the needed mechanical property can be maintained in the early stage of implantation into a body, and the service life of the wire is prolonged. And in the later stage, complete degradation can be gradually realized under the action of implant tissue fluid.

Description

Technical Field

[0001] The present invention belongs to the field of medical biomaterials, and particularly relates to a degradable zinc alloy wire for microtia correction surgery and a preparation method thereof. Background Art

[0002] Biomaterials are materials used to interact with biological systems to treat or enhance tissues, organs, or functions in the body. They can be classified into two categories: bioinert and biodegradable. Biodegradable materials are those that decompose within a certain period after implantation in the body. Biodegradable materials provide specific applications in the body and degrade within a given time once the implantation purpose is completed. Degradation is affected by different variables, including the morphology, composition, and manufacturing process of the materials used, as well as the different implantation sites. Many materials have been developed as biodegradable materials, such as metals, polymers, and ceramics. As an irreplaceable medical device in traditional surgery, sutures have been widely studied. However, the metal sutures used in clinics are often bioinert filaments such as steel wires, titanium, or titanium alloys, which remain in the human body for a long time after surgery. If removed, the patient will experience the pain of a second surgery. If not removed, during daily activities, the inert sutures may move, and their tips are prone to accidents during the free movement process, posing a safety hazard. The development and use of biodegradable sutures avoid the patient experiencing a second surgery and improve the effectiveness and safety of sutures as medical devices.

[0003] Congenital microtia is a common congenital disease, and its treatment mainly focuses on auricular reconstruction. Currently, the most commonly used auricular framework material is autologous costal cartilage. For the preparation of the auricular framework, costal cartilage strips need to be built into a multi-layer structure, and the cartilage in special parts also needs to be bent and shaped. In order to maintain the shape of the auricular framework, steel wires, titanium alloy sutures or synthetic absorbable sutures are often used to splice and fix the costal cartilage strips to obtain a stable three-dimensional structure. Currently, the most commonly used metal suture is titanium alloy. The patent with the authorized publication number CN214549489U invented a titanium-nickel shape memory alloy medical suture, which solves the problem that the suture is easily detached from the wire hole after being inserted into the wire hole by opening different grooves and connecting blocks on the metal wire, and achieves the analgesic effect by storing drugs in the grooves. However, long-term placement of metal sutures in the body will increase the risk of implant-related infections, and titanium alloy sutures may also move during daily activities, and their tips are easy to pierce the skin, resulting in the leakage of titanium wires and cartilage, increasing the risk of postoperative infections. Polymer biodegradable sutures have also been widely used in surgical operations. The patent with the authorized publication number CN117684394A invented a preparation method of an antibacterial and anti-inflammatory high-toughness nylon suture. It uses the method of pre-treating nylon single fibers and then soaking them in an aldehyde-functionalized phenylboronic acid solution and a polyphenol solution to obtain anti-inflammatory high-toughness nylon single fibers, and then weaving them into nylon sutures, and soaking them in an aqueous solution of a cationic polymer and a silver ion-containing compound to obtain an antibacterial and anti-inflammatory high-toughness nylon suture. The patent with the authorized publication number CN117482275A invented a biodegradable medical polymer suture and its preparation method. In the presence of a catalyst, a PLGA-Upy polymer is obtained, immersed in an extracellular matrix solution, and EDC / NHS is added for crosslinking to obtain an ECM-PLGA-Upy composite material, which is processed into a biodegradable suture. Then, through reactions and composites with various solutions, sutures with excellent properties such as antibacterial, controllable degradation rate, good biocompatibility and mechanical strength are prepared. However, synthetic polymers have the accumulation of toxic residual monomers, wear debris and acidic intermediate degradation products, which can lead to pathological bone resorption. In addition, once the surface of synthetic polymer materials is scratched by a suture needle during the suture process, the surface complex will easily fall off, resulting in material failure. Potential risks such as knot slipping and tissue tearing during suture after the suture is used to suture the wound still have not been solved. Poor mechanical properties also limit their further clinical applications. Artificial absorbable sutures have poor stress intensity, cannot continuously fix the cartilage shape, have poor stability, the cartilage framework may gradually deform, and the artificial absorbable sutures degrade slowly, producing acidic metabolites during the degradation process, causing cartilage absorption and surrounding inflammatory reactions. Therefore, it is necessary to develop a biodegradable suture with appropriate mechanical properties to meet the requirements of microtia surgery.

[0004] Biodegradable metals that "biocorrode" have attracted great interest from researchers. Compared with traditional metal implants, biodegradable metals release metal ions during the metal degradation process to promote the repair of surrounding tissues. Compared with synthetic polymers or bioceramics, biodegradable metals have better mechanical properties, including Young's elastic modulus, yield strength, tensile strength and toughness. At present, only magnesium, zinc, iron and their alloys have been reported as biodegradable implant metals. In particular, magnesium-based implants have received attention because of their widely accepted biocompatibility. Magnesium alloys have gradually been studied as degradable suture materials due to their good biocompatibility and degradability. Their natural degradation products can be absorbed or excreted by the human body. Magnesium ions are also basic elements in human metabolism and bone tissue, participating in various biochemical reactions in the body. At the same time, they have multiple biomedical functions such as promoting osteogenesis, promoting vascularization and resisting bacterial infection. The patent with the authorization publication number CN117026040A invented a high-strength and tough Mg-Zn-Zr-Ca magnesium alloy wire with a diameter of 0.6 mm, whose tensile strength is 350 MPa, yield strength is 263 MPa, and elongation is 30%. The patent with the authorization publication number CN116099031A invented a Mg-Y-Nd-Gd-Zr degradable magnesium alloy wire with a diameter of 0.3 mm as a suture material, whose tensile strength is ≥340 MPa, yield strength is ≥292 MPa, and elongation is ≥20%. However, the existing degradable magnesium alloy wire has a low elongation, and is easy to break during the process of adding needles and tying knots when used as surgical sutures. The healing time of microtia surgical wounds is greater than 6 months, and the degradation rate of magnesium alloys is too fast to provide sufficient mechanical support. Iron and its alloys have good formability and stability, high strength, and are ideal candidate materials for high-load tissue implants. However, the biocompatibility of iron corrosion products is controversial, and the human body's acceptance limit for Fe is much lower than that for Zn and Mg (Fe is 10-17 mg / day; Zn is 100-150 mg / day; Mg is 375-500 mg / day). The standard electrode potentials of magnesium, zinc, and iron are -2.372, -0.7618, and -0.447 V, respectively. The high corrosion rate of magnesium and its alloys makes them unable to meet the requirements for high load capacity during tissue healing. The degradation rate of iron is too slow, and it is difficult to completely degrade in the body. The degradation products may remain at the implant site, and do not meet the requirements for degradation time of degradable sutures. The standard electrode potential of zinc is between magnesium and iron, which is more suitable and is expected to meet the requirements for degradation time of sutures.

[0005] Zinc is an essential trace element in the human body, participating in almost all physiological metabolic processes in the human body and affecting the activities of various enzymes in the body. Compared with magnesium, zinc and its alloys do not produce hydrogen gas to form bubbles during the degradation process, which can cause inflammation, tissue necrosis, and hinder the growth and recovery of bones. Zinc is an essential mineral required for normal bone development and bone homeostasis. In addition, zinc has been proven to promote bone recovery. Cu, Mn, and Fe are essential trace elements in the human body. Cu plays an important role in maintaining normal hematopoietic function, promoting connective tissue formation, and maintaining the health of the central nervous system in the human body. Cu also has antibacterial effects. Mn promotes the growth and development of bones and promotes bone synthesis. Fe participates in the synthesis of hemoglobin and the transportation of oxygen. Therefore, the research on using degradable zinc alloy sutures to replace traditional steel wires and titanium wires in ear reconstruction surgery not only shows good clinical application prospects but also can get rid of the embarrassing situation of the market monopoly of titanium alloy sutures and synthetic absorbable sutures by foreign companies. On the other hand, for ear reconstruction surgery, considering the composition and performance requirements of the metal materials of degradable sutures, the designed degradable zinc alloy filaments not only need to control the corrosion rate to match the requirements of the wound healing cycle but also need to meet the requirements of low strength and high plasticity of the sutures during the suture process of ear reconstruction surgery. Summary of the Invention

[0006] An object of the present invention is to overcome the problem of non-degradability of inert metal sutures such as steel wires, titanium, and titanium alloy wires used in current clinical practice, and to overcome the disadvantages of low biocompatibility, high strength, and low plasticity of the reported degradable sutures. The present invention provides a degradable and absorbable zinc alloy suture, which has appropriate strength and high plasticity, meets the requirements of surgical sutures, and innovatively provides a bio-degradable zinc alloy suture with antibacterial effects, a preparation method thereof, and an application thereof on the basis of breaking the technical prejudice that degradable zinc alloy wires are not suitable for use as surgical sutures. After being implanted into the body, the suture can be slowly degraded and absorbed by the human body and will disappear automatically after the surgical wound heals. At the same time, due to its good antibacterial effect, the surgical infection risk can be reduced. The preparation process of the zinc alloy suture is simple, the suture force is strong, the degradation speed is appropriate, and it also has good antibacterial effects. It can replace titanium and titanium alloy sutures used in current clinical applications and obtain better clinical application effects.

[0007] One object of the present invention is to provide a degradable zinc alloy wire for microtia correction surgery, characterized in that the degradable zinc alloy wire is a Zn-Cu-Mn-Fe zinc alloy, and the composition of its alloy elements is calculated by mass percentage as: 0 < Cu ≤ 3 wt.%, 0 < Mn ≤ 1 wt.%, 0 < Fe ≤ 1 wt.%, and the rest is Zn.

[0008] The tensile strength of the degradable zinc alloy wire is 100-150 MPa, the yield strength is 80-120 MPa, the elongation is greater than 58%, preferably greater than 70%, for example, greater than 70% and less than 99%. After soaking in a simulated body fluid (SBF) solution at 37.5 °C for 7 days, the corrosion rate is calculated to be >0.85 mm / y, and after soaking in SBF for 14 days, the corrosion rate is calculated to be >0.4 mm / y.

[0009] The degradable zinc alloy wire is Zn97.6Cu2Mn0.2Fe0.2 or Zn97.4Cu2Mn0.2Fe0.4 or Zn97.4Cu2Mn0.4Fe0.2; preferably, the degradable zinc alloy wire is Zn97.6Cu2Mn0.2Fe0.2. Further, when made into a suture with a diameter of 0.2 mm, its tensile strength is 125.2 MPa and the elongation is 78.7%.

[0010] The second object of the present invention is to provide a preparation method of a degradable zinc alloy wire, which is characterized in that the degradable zinc alloy wire is a Zn-Cu-Mn-Fe zinc alloy, and the alloy element composition is by mass percentage: 0 < Cu ≤ 3 wt.%, 0 < Mn ≤ 1 wt.%, 0 < Fe ≤ 1 wt.%, and the rest is Zn; the preparation method includes the following steps:

[0011] (1) Weigh high-purity raw materials of Zn, Cu, Mn, and Fe, and melt them under vacuum and protective atmosphere, cast them, and obtain a zinc alloy ingot after natural cooling;

[0012] (2) Homogenize the zinc alloy ingot and then perform backward extrusion to obtain a zinc alloy bar;

[0013] (3) Perform room-temperature cold drawing processing and on-line annealing hot drawing on the zinc alloy bar to obtain a zinc alloy wire.

[0014] According to one embodiment, step (1) includes: preparing raw materials according to the components and contents of the zinc alloy, placing the raw materials in a crucible, melting and refining them fully in a vacuum induction furnace under argon protective atmosphere, and then casting. The melting and pouring temperature is 500-600 °C;

[0015] According to one embodiment, step (2) includes: placing the ingot obtained in step (1) in a medium-temperature furnace, holding it at 200-400 °C for 24-48 h for homogenization, then preheating the ingot and the mold to 180-300 °C, holding for 30-60 minutes, and performing backward extrusion. The extrusion speed is 1-5 mm / s, the extrusion ratio is 50-100, preferably 70-90, to obtain a zinc alloy bar with a diameter of 4-10 mm;

[0016] According to one embodiment, step (3) includes: when the diameter of the bar obtained in step (2) is greater than 1 mm, cold drawing is performed at a speed of 1-10 mm / s, preferably 1-5 mm / s, with a single-pass deformation amount of 8-20%; when the diameter is less than 1 mm, hot drawing is performed at a speed of 1-10 mm / s, preferably 1-5 mm / s, with a single-pass deformation amount of 5-15%, and the temperature is 200-300 °C; preferably, cold drawing also includes annealing at 300-330 °C for 20-40 min when the cumulative deformation amount is 45-100%.

[0017] According to one embodiment, the diameter range of the zinc alloy wire in step (3) is 0.1-0.6 mm, preferably 0.2-0.3 mm. In one embodiment, the diameter range of the zinc alloy wire is 0.1 mm, 0.15 mm, 0.2 mm, 0.25 mm, 0.3 mm, 0.35 mm, 0.4 mm, 0.45 mm, 0.5 mm, 0.55 mm, or 0.6 mm.

[0018] The third object of the present invention is a degradable zinc alloy wire obtained by the foregoing preparation method.

[0019] The fourth object of the present invention is to provide the use of the foregoing degradable zinc alloy wire in making medical sutures; preferably, the medical sutures are applied to surgeries such as plastic surgery, urology, pediatrics, stomatology, otolaryngology, orthopedics, sports medicine, etc. and the suture of skin and soft tissues, especially microtia correction surgery.

[0020] The fifth object of the present invention is to provide a finished suture product, including a suture and a suture needle, wherein the suture is made of the foregoing degradable zinc alloy wire.

[0021] Beneficial effects:

[0022] (1) The present invention adopts a backward extrusion method with a large extrusion ratio to effectively improve the mechanical properties of zinc alloy wires. Compared with forward extrusion, there is no friction between the ingot and the extrusion cylinder in backward extrusion, and the required extrusion force is much smaller than that in forward extrusion. Moreover, compared with forward extrusion, the extruded bar obtained by backward extrusion can effectively refine the grains, balance the material fluidity, eliminate defects such as pores and looseness in the tissue, and improve the mechanical properties and corrosion uniformity of the material. Through a large extrusion ratio, the grains are better refined, internal defects are eliminated, and the degree of internal grain refinement can be controlled by adjusting the extrusion ratio, so that the grains are refined to a certain size to meet the material requirements for subsequent cold and hot drawing.

[0023] (2) The present invention adopts a process of first cold drawing and then hot drawing for alloy wires of different diameters. When the diameter of the bar is relatively thick, the influence of minor defects on the drawing process is relatively small, and when the bar is thicker, its length is also shorter. It is difficult to directly perform hot drawing after heating through a tube furnace. Thicker bars require longer annealing time. By annealing to remove stress first and then performing cold drawing, the drawing time can be shortened. When the diameter of the wire is less than 1 mm, defects have a greater impact on the properties of the material, and at this time, the length of the wire can pass through the tube furnace through previous drawing. By real-time annealing at a lower temperature, certain internal stress can be eliminated, the structure can be made uniform, the properties of the wire can be improved, and the occurrence of wire breakage during the drawing process can be reduced. Hot drawing can also reduce the time for removing the material for annealing, improve the efficiency of the drawing process. The alloy wire prepared by hot drawing has better surface quality and higher elongation, making the wire have better corrosion resistance and knotting performance. By the drawing process of first cold drawing combined with subsequent hot drawing, the risk of generating microcracks during the drawing process can be effectively reduced and the working efficiency can be improved.

[0024] (3) The prepared medical zinc alloy suture of the present invention adopts a unique composition design. Medical materials are required to have high biocompatibility. Cu, Mn, and Fe elements are all elements with high safety. The Cu element can enhance the plasticity and antibacterial properties of the zinc alloy wire, and has good biological and cell compatibility, which can effectively avoid unnecessary consequences caused by bacterial infection during the application of the suture. The Mn element can greatly improve the plasticity of the zinc alloy. Mn can promote the proliferation, adhesion, and diffusion of osteoblasts, and has good biocompatibility. The addition of Mn is expected to optimize the material properties and improve the biocompatibility of the Zn alloy. The Fe element is an essential nutrient element in the human body and is also a component of hemoglobin. Pure iron has been proven to have good biocompatibility. The addition of the Fe element can promote the uniform corrosion of the zinc alloy, reducing the possibility of premature failure of the suture due to local preferential fracture during the corrosion process. For the alloy wire prepared from this alloy composition, due to the addition of Cu, Fe, and Mn elements, CuZn4 phase and (Mn, Fe)Zn13 phase are formed. The CuZn4 phase is elongated and broken during the multi-pass drawing process with large deformation and is evenly distributed along the drawing direction, promoting the occurrence of uniform corrosion. The (Mn, Fe)Zn13 belongs to the hard phase. Although it is fragmented to a certain extent along the drawing direction during the drawing process, it still partially retains a relatively large size. There is a large potential difference between this phase and the matrix, promoting the occurrence of rapid corrosion, and this phase is widely distributed throughout the alloy wire. In summary, this alloy wire has a relatively high corrosion rate and a relatively uniform corrosion morphology.

[0025] (4) The prepared medical zinc alloy suture of the present invention has a higher elongation compared with other existing zinc alloy wires.

[0026] The alloy wire exhibits a high elongation rate, which is mainly attributed to the combined effect of several key factors. First, the large deformation in the drawing process optimizes the microstructure of the material. The large plastic deformation results in a significant reduction in grain size, forming a fine and uniform grain structure. This fine grain structure not only helps to improve the strength of the material, but also the increased grain boundaries can prevent the propagation of cracks, increase the crack propagation path, and benefit the plasticity of the material. Second, the combined use of cold and hot drawing processes plays a crucial role in the production of alloy wires. During the cold drawing process, the material is stretched at room temperature, which helps to obtain higher strength and hardness, while hot drawing is carried out above the recrystallization temperature of the material, which helps to release and reduce internal stresses and prevent the hardening and embrittlement of the material. This process with temperature change not only promotes the recrystallization process inside the material, but also helps to obtain finer and more uniform grains, while reducing the internal stresses generated due to work hardening. Finally, the second phase in the alloy wire is distributed in the matrix in a fine and uniform manner, avoiding stress concentration, thus maintaining good plasticity and toughness of the material.

[0027] In summary, the high elongation rate of this alloy wire is due to the combined effects of grain refinement during the drawing process, stress control in the cold and hot drawing processes, and the uniform distribution of the second-phase particles.

[0028] (5) The medical zinc alloy suture prepared by the present invention has a more uniform corrosion rate compared with other existing zinc alloy wires. A significant defect of the existing zinc alloy wires as sutures is that their degradation rate is too slow. Compared with the Zn-Cu-Mn alloy, the Zn-Cu-Mn-Fe alloy wire developed in this study by adding an appropriate amount of the nutrient element Fe has a more uniform corrosion morphology and a higher corrosion rate. The addition of Fe promotes the precipitation of massive Fe-rich phases, and there is a large potential difference between this phase and the matrix, which promotes the occurrence of galvanic corrosion, increases the degradation rate, and the Fe-rich phases are widely distributed along the drawing direction, which makes the occurrence of galvanic corrosion spread throughout the entire alloy wire simultaneously, thus promoting uniform corrosion. And by observing the microstructure of the extruded state, it is found that the addition of Fe has the effect of refining the size of the precipitated phases and promoting the uniform distribution of the precipitated phases. Brief Description of the Drawings

[0029] Figure 1 is the prepared Zn-2Cu-0.2Mn-0.2Fe alloy suture.

[0030] Figure 2 is the stress-strain curve of the prepared Zn-2Cu-XMn-XFe alloy suture.

[0031] Figure 3 is the schematic diagram of the prepared Zn-2Cu-0.2Mn-0.2Fe alloy suture after adding the needle.

[0032] Figure 4 Polarization curves of the prepared Zn-2Cu-0.2Mn-0.2Fe alloy suture in simulated body fluid (SBF).

[0033] Figure 5 Microstructure of the prepared Zn-2Cu-0.2Mn-0.2Fe alloy suture, with two evenly and densely distributed precipitation phases in the alloy wire.

[0034] Figure 6 Comparison chart of corrosion rates at different times of the zinc alloy wire of Example 1 and the zinc alloy wire of the comparative example after soaking in SBF solution.

[0035] Figure 7 Surface morphology diagrams of the zinc alloy wire of Example 1 and the zinc alloy wire of the comparative example after removing corrosion products after soaking in SBF solution for 14 days, where (a) is the surface morphology diagram of the Zn-2Cu-0.2Mn-0.2Fe alloy wire after removing corrosion products after soaking in SBF solution for 14 days; (b) is the surface morphology diagram of the Zn-2Cu-0.2Mn alloy wire after removing corrosion products after soaking in SBF solution for 14 days.

[0036] Figure 8 Cytotoxicity test results of the Zn-2Cu-0.2Mn-0.2Fe alloy suture of Example 1.

[0037] Figure 9 Blood compatibility experiment of the Zn-2Cu-0.2Mn-0.2Fe alloy suture of Example 1.

[0038] Figure 10 Antibacterial performance experiment results of the Zn-2Cu-0.2Mn-0.2Fe alloy suture of Example 1. Specific implementation manners

[0039] Next, the specific implementation manners of the present invention will be described in detail with reference to the accompanying drawings.

[0040] Next, the technical solutions and the technical problems to be solved in the embodiments of the present invention will be elaborated. Obviously, the described embodiments are only a part of the embodiments of the present invention patent, rather than all embodiments. Without departing from the concept of the present invention, several adjustments and improvements can be made, and all of these belong to the patent protection scope of the present invention.

[0041] The present invention provides a degradable zinc alloy wire especially for microtia correction surgery. The degradable zinc alloy wire is a Zn-Cu-Mn-Fe zinc alloy, and its alloy element composition is by mass percentage: 0 < Cu ≤ 3 wt.%, 0 < Mn ≤ 1 wt.%, 0 < Fe ≤ 1 wt.%, and the rest is Zn.

[0042] The tensile strength of the degradable zinc alloy wire is 100 - 150 MPa, the yield strength is 80 - 120 MPa, the elongation is greater than 58%, preferably greater than 75%. After soaking in SBF solution at 37.5 °C for 7 days, its corrosion rate is calculated to be > 0.85 mm / y, and after soaking in SBF solution for 14 days, its corrosion rate is calculated to be > 0.4 mm / y. The Zn-Cu-Mn-Fe quaternary alloy wire of the present invention has excellent plasticity, suitable tensile strength and yield strength required for the suture wire in microtia correction surgery, can maintain the required mechanical properties in the initial stage of implantation in the body, and can be gradually completely degraded under the action of the implant tissue fluid in the later stage.

[0043] The degradable zinc alloy wire is Zn97.6Cu2Mn0.2Fe0.2 or Zn97.4Cu2Mn0.2Fe0.4 or Zn97.4Cu2Mn0.4Fe0.2; preferably, the degradable zinc alloy wire is Zn97.6Cu2Mn0.2Fe0.2. Further, when made into a suture with a diameter of 0.2 mm, its tensile strength is 125.2 MPa and the elongation is 78.7%.

[0044] The present invention provides a preparation method of a degradable zinc alloy wire especially for microtia correction surgery, and its preparation includes the following steps:

[0045] (1) Vacuum induction melting of pure zinc ingots and master alloys under the protection of an inert atmosphere. The master alloys are Cu-Zn master alloy, Zn-Mn master alloy, and Zn-Fe master alloy. After full melting, pouring is carried out under a protective atmosphere to obtain a zinc alloy ingot. The zinc alloy ingot is a Zn-Cu-Mn-Fe quaternary alloy, and the mass percentage of each element is: the content of Cu is 0% - 3 wt.%, the content of Mn is 0% - 1 wt.%, the content of Fe is 0% - 1 wt.%, and the rest is Zn.

[0046] (2) Homogenization annealing of the above zinc alloy ingot is carried out to ensure uniform structure and provide guarantee for subsequent large deformation.

[0047] (3) Cylindrical blanks are cut from the above zinc alloy ingot, and backward extrusion processing with a large extrusion ratio is carried out to obtain zinc alloy bars with fine and uniform grains. The diameter of the bars is 4 - 10 mm.

[0048] (4) The zinc alloy bar is drawn to obtain the degradable zinc alloy wire. The degradable zinc alloy wire has low strength, high toughness, is easy to suture and deform, and has a certain mechanical stability.

[0049] Preferably, in step (1), the melting process is as follows: High-purity zinc and intermediate alloy ingots are added to a melting furnace, the vacuum is pumped to below 0.1 Pa, and then the protective gas argon is filled. The temperature is raised to 500 - 650 °C, held for 15 - 30 min, then cooled to 300 - 400 °C and held for 20 min for casting. The final size of the cast bar is 80 mm in diameter and 200 mm in height.

[0050] Preferably, in step (2), the homogenization process is as follows: The temperature of the medium-temperature furnace is raised to 200 - 400 °C, the cast bar from step (1) is put in, held for 24 - 48 h, and then the cast bar is taken out and air-cooled to room temperature.

[0051] Preferably, in step (3), the extrusion process is as follows: The cast bar is processed into a cylinder with a diameter of 35 - 40 mm and a height of 15 - 30 mm, held at 180 - 300 °C for 30 - 60 minutes, and backward extruded at a speed of 1 - 5 mm / s. The extrusion ratio is 20 - 100, and a bar with a diameter of 4 - 10 mm is backward extruded.

[0052] Preferably, in step (4), the room-temperature drawing process is as follows: The alloy bar obtained by backward extrusion in step (3) is machined to make the alloy bar straight, smooth and flat. After annealing at an annealing temperature of 200 - 400 °C for 30 - 60 min, it is cooled with the furnace. Then room-temperature cold drawing is carried out. The drawing speed is 1 - 10 m / min, preferably 1 - 5 mm / s. The single-pass deformation amount is 8% - 20%. When the cumulative deformation amount reaches 45% - 100%, intermediate annealing is carried out at an annealing temperature of 200 - 400 °C. This process is repeated until the diameter of the alloy bar is 1 mm. To prevent wire breakage when the diameter is less than 1 mm, hot drawing is then carried out at a drawing temperature of 200 - 300 °C and a drawing speed of 1 - 10 m / min, with a pass deformation amount of 5% - 15%. The alloy wire is passed through a tube-type heating furnace, and the temperature in the heating furnace is set at 200 - 400 °C, and the temperature in the furnace is detected in real time by a thermocouple.

[0053] A finished suture product includes a suture and a suture needle, and the suture is made of the aforementioned degradable zinc alloy wire.

[0054] The preparation process of the finished suture is as follows: Select a suture needle made of high-strength stainless steel material. There are tiny holes designed at the tail of the needle. The diameter of the holes is slightly larger than the diameter of the suture, for example, 0.3 mm, and the shape is circular. The zinc alloy suture is treated with ultraviolet sterilization to ensure a sterile state. Use a high-precision crimping device equipped with a numerical control system that can precisely control the crimping position and force. The device includes a high-strength crimping head and a positioning device to ensure the accuracy of the crimping process. The crimping device is integrated with a real-time quality monitoring system that can detect pressure, position, and time parameters during the crimping process to ensure the stability and consistency of each crimping. First, perform positioning. Insert one end of the suture into the hole at the tail of the needle, and the insertion depth is 1 mm. Use a high-power magnifying glass for operation to ensure that the wire is horizontally aligned with the hole. Then, perform preliminary fixation to prevent movement or deviation during the wire insertion process. After that, perform crimping. The robotic arm of the crimping device precisely positions the suture needle under the crimping head to ensure that the hole at the tail of the needle is completely aligned with the crimping head. Set the crimping parameters, including pressure, time, and the shape of the crimping head. The shape of the crimping head used is cylindrical. Start the crimping device, and the crimping head crimps the tail of the needle with the set pressure and time. The material of the crimping head is high-hardness alloy, which applies uniform pressure within a very small range to ensure that the wire and the metal at the tail of the needle undergo plastic deformation to achieve a firm mechanical connection. Finally, perform quality inspection. Use a microscope to visually inspect the crimping part to ensure that the connection between the wire and the tail of the needle is firm, without obvious damage or looseness. Conduct a tensile test. The test device applies gradually increasing tensile force until a tensile force value of 15 N is reached to ensure that the wire will not fall off during the operation. Then, conduct a metallographic analysis of the crimping part to observe the bonding situation between the wire and the metal at the tail of the needle to ensure that there are no microcracks or defects. The finished product is packaged aseptically. The packaging material is an aseptic paper bag to ensure that it is not contaminated during transportation and storage. The packaged suture needle is subjected to secondary radiation sterilization and disinfection treatment.

[0055] The applications of the above-mentioned biodegradable and absorbable zinc alloy suture include surgeries in plastic surgery, urology, pediatrics, stomatology, otolaryngology, orthopedics, sports medicine, etc., and the suturing of dermal soft tissues, especially for the suturing in ear reconstruction surgery.

[0056] Example 1: Zn-2Cu-0.2Mn-0.2Fe

[0057] (1) Weigh the required alloy components according to the weight percentages of 2 wt.% Cu, 0.2 wt.% Mn, 0.2 wt.% Fe, and the balance being Zn. The purity of the zinc raw material is greater than or equal to 99.99%, the purity of the Cu-30Zn master alloy is greater than or equal to 99.99%, the purity of the Zn-20Mn master alloy is greater than or equal to 99.99%, and the purity of the Zn-2Fe master alloy is greater than or equal to 99.99%.

[0058] (2) After the surface of the raw material is polished clean, put the above-mentioned high-purity metal into the crucible of a vacuum melting furnace, evacuate to 0.1 Pa, then fill with 10,000 Pa argon, start heating until the metal is completely melted, and after 20 minutes of sufficient refining, pour it into an ingot of specified size. The melting and pouring temperature is 500 - 600 °C. After cooling, break the vacuum and take out the ingot for use. The final ingot size is 80 mm in diameter and 200 mm in height.

[0059] (3) Subsequently, place the ingot in a medium-temperature furnace for homogenization treatment. The holding temperature is 400 °C, the holding time is 36 h, and after holding, air-cool to room temperature.

[0060] (4) Process the homogenized ingot into a cylindrical ingot with a diameter of 40 mm and a length of 20 mm. Hold at 180 - 300 °C for 30 - 60 minutes, then perform backward extrusion processing. The extrusion temperature is 300 °C, the extrusion speed is 5 mm / s, the extrusion ratio is 80, and backward extrude a bar with a diameter of 4.5 mm. After machining and polishing the surface of the sample, perform wire drawing.

[0061] (5) Perform annealing at 300 °C for 30 min before wire drawing. The wire drawing speed is 2 m / min, and the wire drawing temperature is room temperature. The single-pass deformation is 10%. When the cumulative deformation is about 50%, perform annealing at 300 °C for 30 min, and repeat this process until the wire diameter is 1 mm. Pass the wire through a tube furnace, set the furnace temperature to 300 °C, keep the wire in the furnace for 10 min and then perform hot wire drawing. The wire drawing speed is 2 m / min, and the single-pass deformation is 10% until the wire diameter is 0.2 mm. Figure 1 The macroscopic diagram of the prepared Zn-2Cu-0.2Mn-0.2Fe biodegradable absorbable magnesium alloy suture with a diameter of 0.2 mm. By testing the Zn-2Cu-0.2Mn-0.2Fe alloy obtained through the above steps, the tensile strength at room temperature is ≧ 100 Mpa, for example, the tensile strength is 125.2 MPa, the yield strength is ≧ 80 MPa, and the elongation is ≧ 75%, for example, the elongation is 78.7%. As Figure 2 shown, the material properties meet the requirements for the preparation of medical sutures. As Figure 3 shown, add needles to the alloy wire to make the finished suture.

[0062] The corrosion performance was tested. Figure 4 The polarization curve of the prepared Zn-2Cu-0.2Mn-0.2Fe wire in SBF simulated body fluid at 37.5 °C. The corrosion potential Ecorr of the alloy wire is -1.14 V, and the corrosion current Icorr is 5.15×10 -5The alloy wire was immersed in SBF simulated body fluid, and its corrosion rate was 0.42 mm / y. The alloy wire corroded uniformly. Its microstructure is as Figure 5 shown. The left figure is the microstructure under the optical microscope, and the right figure is the microstructure under the scanning electron microscope.

[0063] Biological evaluation of Zn-2Cu-0.2Mn-0.2Fe

[0064] A. Cytotoxicity evaluation

[0065] Before preparing the leaching solution, the wire materials were ultrasonically cleaned with absolute ethanol, disinfected with 75% alcohol, and then the residual ethanol was washed with the culture medium. Then, a 0.2 g / ml leaching solution was prepared with the complete culture medium and incubated for 24 h. The cells were seeded in a 48-well plate at a density of 5×10 3 cells / well and allowed to adhere for 24 h. Then, the culture medium was replaced with the leaching solution prepared from the wire materials (the concentrations of the leaching solution were 10%, 40%, 50%, and 100% respectively). The control group was added with normal culture medium. Incubate for 24 h, add the complete culture medium containing 10% CCK-8 solution to each well, and culture in a humidified incubator at 37°C and 5% CO 2 for 2 hours. Then, transfer 100 μL of the culture medium from each well to a 96-well plate, and use an enzyme-linked immunosorbent assay (ELISA) reader to read the results at an emission wavelength of 450 nm to measure the absorbance values of each group. As Figure 8 shown, the results indicate that there was no significant difference in the absorbance values of the leaching solutions at different concentrations compared with the control group, proving that the wire material has good biocompatibility with cells.

[0066] Hemolysis experiment evaluation

[0067] Take 500 μl of whole blood, add 5 ml of PBS solution, centrifuge at 5000 rpm for 5 min, collect the precipitate and add PBS solution to 10 ml, centrifuge at 5000 rpm for 5 min, collect the precipitate again and add PBS solution to 10 ml, centrifuge at 5000 rpm for 5 min until the supernatant is clear, and then add 10 ml of PBS solution to resuspend the live cells. Positive control group: Take 200 μl of red blood cell suspension + 800 μl of ddH2O, negative control group take 200 μl of red blood cell suspension + 800 μl of PBS solution, experimental group: take 200 μl of red blood cell suspension + 800 μl of PBS leaching solution prepared at different concentrations (the concentrations of the leaching solution are 10%, 40%, 50%, 70%, and 100% respectively), incubate at 37°C for 1 h, and then centrifuge at 5000 rpm for 5 min to observe and compare the hemolysis reaction. The observation results show ( Figure 9 ), the hemolysis of red blood cells treated with leaching solutions of different concentrations of wire materials was significantly lower than that of the pure water treatment group, and even the 100% concentration leaching solution showed no obvious hemolysis phenomenon, proving that the wire material can avoid hemolysis.

[0068] Evaluation of antibacterial effect

[0069] Take the colonies of Staphylococcus aureus, dilute them to 0.3 McFarland turbidity with PBS solution, then take 5 ml of the bacterial suspension, add the wire material to the experimental group, and the volume ratio of the wire material to PBS is 10 mg / ml. The control group is not treated. Incubate at 125 rpm in a shaker at 37 °C for 24 h. Then take 40 μl of the bacterial suspension, spread it on an agar plate containing bacterial medium, and then incubate the agar plate in an incubator at 37 °C for 24 h. Observe the growth of bacterial colonies. The antibacterial results show ( Figure 10 ), in the group co-incubated with the wire material, the number of CFU plated bacterial colonies is significantly less than that of the control group, proving that the wire material of the present invention has antibacterial advantages.

[0070] Example 2: Zn-2Cu-0.2Mn-0.4Fe

[0071] (1) Weigh the required alloy components according to the weight percentages of 2 wt.% Cu, 0.2 wt.% Mn, 0.4 wt.% Fe, and the balance being Zn, where the purity of the zinc raw material is greater than or equal to 99.99%, the purity of the Cu-30Zn master alloy is greater than or equal to 99.99%, the purity of the Zn-20Mn master alloy is greater than or equal to 99.99%, and the purity of the Zn-2Fe master alloy is greater than or equal to 99.99%.

[0072] (2) After the surfaces of the raw materials are polished clean, put the above high-purity metals into the crucible of a vacuum melting furnace, evacuate to 0.1 Pa, then fill with 10,000 Pa of argon, start heating until all the metals are melted, and after 20 minutes of sufficient refining, pour into ingots of specified dimensions. After cooling, take out the ingots under reduced pressure for use. The final ingot size is 80 mm in diameter and 200 mm in height.

[0073] (3) Subsequently, perform homogenization treatment on the ingots, with the holding temperature of 380 °C and the holding time of 36 h, and then air-cool to room temperature after holding.

[0074] (4) Process the homogenized ingots into cylindrical ingots with a diameter of 40 mm and a length of 20 mm, hold at a temperature between 180 - 300 °C for 30 - 60 minutes, then perform backward extrusion processing, with the extrusion temperature of 300 °C, the extrusion speed of 4 mm / s, the extrusion ratio of 64, and backward extrude bars with a diameter of 5 mm. After machining and polishing the surface of the samples, perform drawing.

[0075] (5) Annealing is carried out at 320 °C for 30 min before drawing, the drawing speed is 2 m / min, and the drawing temperature is room temperature. The single-pass deformation is 10%, and annealing at 320 °C for 30 min is carried out when the cumulative deformation is 45%. Repeat this process until the wire diameter is 1 mm. Pass the wire through a tube furnace, set the furnace temperature to 320 °C, keep the wire in the furnace for 10 min and then carry out hot drawing, the drawing speed is 2 m / min, and the single-pass deformation is 10% until the wire diameter is 0.2 mm.

[0076] Testing the Zn-2Cu-0.2Mn-0.4Fe alloy obtained by the above steps, the tensile strength at room temperature is 133 MPa, and the elongation is 58.28%, as Figure 2 shown. Carry out suture needle addition on the alloy wire, and the material properties also meet the requirements for the preparation of medical sutures.

[0077] (6) The corrosion performance was tested. The alloy wire was immersed in SBF simulated body fluid, and the alloy wire corroded uniformly, and its corrosion rate was 0.65 mm / y.

[0078] Example 3: Zn-2Cu-0.4Mn-0.2Fe

[0079] (1) Weigh the required alloy components according to the weight percentages of 2 wt.% Cu, 0.4 wt.% Mn, 0.2 wt.% Fe, and the balance being Zn, where the purity of the zinc raw material is greater than or equal to 99.99%, the purity of the Cu-30Zn master alloy is greater than or equal to 99.99%, the purity of the Zn-20Mn master alloy is greater than or equal to 99.99%, and the purity of the Zn-2Fe master alloy is greater than or equal to 99.99%.

[0080] (2) After the surfaces of the raw materials are polished clean, put the above high-purity metals into the crucible of a vacuum melting furnace, evacuate to 0.1 Pa, then fill with 10,000 Pa argon, turn on the heating until all the metals are melted, and after 25 minutes of full refining, pour into ingots of specified size. After cooling, take out the ingots under reduced pressure for use. The final ingot size is 80 mm in diameter and 200 mm in height.

[0081] (3) Subsequently, carry out homogenization treatment on the ingots, the holding temperature is 400 °C, the holding time is 36 h, and after holding, air-cool to room temperature.

[0082] (4) Process the homogenized ingots into cylindrical ingots with a diameter of 40 mm and a length of 20 mm, hold for 30 - 60 minutes between 180 - 300 °C, and then carry out backward extrusion processing. The extrusion temperature is 300 °C, the extrusion speed is 3 mm / s, the extrusion ratio is 64, and backward extrude a rod with a diameter of 5 mm. After machining and grinding the surface of the sample, carry out drawing.

[0083] (5) Before drawing, annealing is carried out at 320 °C for 30 min, the drawing speed is 2 m / min, and the drawing temperature is room temperature. The single-pass deformation is 10%, and annealing at 320 °C for 30 min is carried out when the cumulative deformation is 55%. Repeat this process until the wire diameter is 1 mm. Pass the wire through a tube furnace, set the furnace temperature to 320 °C, keep the wire in the furnace for 10 min and then carry out hot drawing, the drawing speed is 2 m / min, and the single-pass deformation is 10% until the wire diameter is 0.2 mm.

[0084] The Zn-2Cu-0.2Mn-0.4Fe alloy obtained through the above steps is tested, and the tensile strength at room temperature is 129.08 MPa and the elongation is 93.1%, as Figure 2 shown. Carry out suture needle adding on the alloy wire, and the material properties meet the requirements for the preparation of medical suture.

[0085] (6) The corrosion performance is tested. Immerse the alloy wire in SBF simulated body fluid, and the alloy wire corrodes uniformly, and its corrosion rate is 0.58 mm / y.

[0086] Comparative example: Zn-2Cu-0.2Mn

[0087] (1) In this comparative example, the required alloy components are weighed according to the weight percentages of 2 wt.% Cu, 0.2 wt.% Mn, and the balance being Zn. The purity of the zinc raw material is greater than or equal to 99.99%, the purity of the Cu-30Zn master alloy is greater than or equal to 99.99%, and the purity of the Zn-20Mn master alloy is greater than or equal to 99.99%.

[0088] (2) After the surfaces of the raw materials are polished clean, put the above high-purity metals into the crucible of a vacuum melting furnace, evacuate to 0.1 Pa, then fill with 10,000 Pa argon, start heating until all the metals are melted, and after 20 minutes of sufficient refining, pour into ingots of specified dimensions. After cooling, take out the ingots under reduced pressure for use. The final ingot size is 80 mm in diameter and 200 mm in height.

[0089] (3) Subsequently, the ingots are subjected to homogenization treatment, the holding temperature is 400 °C, the holding time is 36 h, and after holding, air-cool to room temperature.

[0090] (4) Process the homogenized ingots into cylindrical ingots with a diameter of 40 mm and a length of 20 mm, keep them at 180 - 300 °C for 30 - 60 minutes, and then carry out backward extrusion processing. The extrusion temperature is 300 °C, the extrusion speed is 5 mm / s, the extrusion ratio is 80, and backward extrude a bar with a diameter of 4.5 mm. After machining and polishing the surface of the sample, carry out drawing.

[0091] (5) Annealing was carried out at 320 °C for 30 min before drawing. The drawing speed was 2 m / min, and the drawing temperature was room temperature. The single-pass deformation was 10%, and annealing at 300 °C for 30 min was carried out when the cumulative deformation was about 50%. This process was repeated until the wire diameter was 1 mm. The wire was passed through a tube furnace, the temperature in the furnace was set at 300 °C, and hot drawing was carried out after the wire was kept warm in the furnace for 10 min. The drawing speed was 2 m / min, and the single-pass deformation was 10% until the wire diameter was 0.2 mm.

[0092] The Zn-2Cu-0.2Mn alloy obtained through the above steps was tested, and the tensile strength at room temperature was 190.27 MPa and the elongation was 87.15%, as Figure 2 shown.

[0093] (6) The corrosion performance was tested. The alloy wire was immersed in SBF simulated body fluid. The corrosion of the alloy wire was uneven, and its corrosion rate was 0.29 mm / y.

[0094] Comparison of corrosion rate and uniform corrosion degree

[0095] The corrosion rate tests were carried out on the comparative example and the example. As Figure 6 shown is the comparison of the corrosion rates of Example 1 and the comparative example. The comparison shows that the corrosion rate of the zinc alloy suture prepared in the example of the present invention is faster. Figure 7 Figure 18 is the surface morphology diagram after removing the corrosion products after Example 1 and the comparative example were immersed in the SBF solution for 14 days. The degradation morphologies of the example and the comparative example were observed. It can be seen that the corrosion morphology of the example is more uniform, the corrosion degree is more serious, and it better meets the requirements of the suture for microtia surgery for a faster corrosion rate and a uniform corrosion degree. At the same time, the biological evaluation shows that it has good cell safety and anti-hemolysis reaction and has the advantage of antibacterial.

[0096] The above are only the preferred embodiments of the present invention and are not used to limit the present invention. For those skilled in the art, the present invention can have various changes and modifications. Any modification, equivalent replacement, improvement, etc. made within the spirit and principle of the present invention shall be included within the scope of the patent protection of the present invention.

Claims

1. A degradable zinc alloy wire for microtia correction, characterized in that: The degradable zinc alloy wire is a Zn-Cu-Mn-Fe zinc alloy, and the alloy element composition thereof is as follows in mass percentage: 0<Cu≤3wt.%, 0<Mn≤1wt.%, 0<Fe≤1wt.%, and the rest is Zn.

2. The degradable zinc alloy wire according to claim 1, characterized in that: The degradable zinc alloy wire has a tensile strength of 100-150 MPa, a yield strength of 80-120 MPa, an elongation greater than 58%, preferably greater than 70%, and a corrosion rate calculated to be >0.85 mm / y after immersion in an SBF solution at 37.5°C for 7 days, and a corrosion rate calculated to be >0.4 mm / y after immersion in an SBF solution for 14 days.

3. The degradable zinc alloy wire according to claim 1, characterized in that: The degradable zinc alloy wire is Zn97.6Cu2Mn0.2Fe0.2 or Zn97.4Cu2Mn0.2Fe0.4 or Zn97.4Cu2Mn0.4Fe0.2; Preferably, the degradable zinc alloy wire is Zn97.6Cu2Mn0.2Fe0.

2. Furthermore, when it is prepared into a suture with a diameter of 0.2 mm, its tensile strength is 125.2 MPa and its elongation is 78.7%.

4. A method for preparing a degradable zinc alloy wire, characterized in that: The degradable zinc alloy wire is a Zn-Cu-Mn-Fe zinc alloy, and the alloy element composition thereof is as follows in terms of mass percentage: 0<Cu≤3wt.%, 0<Mn≤1wt.%, 0<Fe≤1wt.%, and the rest is Zn; the preparation method comprises the following steps: (1) weighing high-purity raw materials of Zn, Cu, Mn, and Fe, and smelting them under vacuum and protective atmosphere, casting them, and naturally cooling them to obtain a zinc alloy ingot; (2) homogenizing the zinc alloy ingot and then performing reverse extrusion to obtain a zinc alloy rod; (3) The zinc alloy rod is subjected to room temperature cold drawing combined with online annealing hot drawing to obtain a zinc alloy wire.

5. The method for preparing the degradable zinc alloy wire according to claim 4, characterized in that: The step (1) comprises: preparing raw materials according to the composition and content of the zinc alloy, placing the raw materials in a crucible, and casting after fully melting and refining in a vacuum induction furnace under an argon protective atmosphere, wherein the melting and pouring temperature is 500-600° C.; Preferably, the step (2) comprises: placing the ingot obtained in the step (1) in a medium temperature furnace, keeping it at 200-400° C. for 24-48 hours for homogenization, then preheating the ingot and the mold to 180-300° C., keeping it at this temperature for 30-60 minutes, and performing reverse extrusion, the extrusion speed is 1-5 mm / s, the extrusion ratio is 50-100, preferably 70-90, to obtain a zinc alloy rod with a diameter of 4-10 mm; Preferably, the step (3) comprises: when the diameter of the rod obtained in the step (2) is greater than 1 mm, cold drawing is performed at a speed of 1-10 mm / s, preferably 1-5 mm / s, and the deformation amount of a single pass is 8-20%; when the diameter is less than 1 mm, hot drawing is performed at a speed of 1-10 mm / s, preferably 1-5 mm / s, and the deformation amount of a single pass is 5-15%, and the temperature is 200-300°C; preferably, the cold drawing also includes annealing at 300-330°C / 20-40min when the cumulative deformation amount is 45-100%.

6. The method for preparing the degradable zinc alloy wire according to claim 4, characterized in that: The diameter of the zinc alloy wire in step (3) is in the range of 0.1-0.6 mm, preferably 0.2-0.3 mm.

7. The degradable zinc alloy wire obtained according to the preparation method according to any one of claims 4 to 6.

8. Use of the degradable zinc alloy wire according to claim 7 in making medical sutures; preferably, the medical sutures are used in plastic surgery, urology, pediatrics, dentistry, otolaryngology, orthopedics, sports medicine and other surgeries and suturing of intracutaneous soft tissues, especially microtia correction.

9. A finished suture product, comprising a suture and a suture needle, wherein the suture is made of the degradable zinc alloy wire material according to claims 1 to 3 or claim 7.

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