Low-cost and high-supporting-performance nickel-titanium shape memory alloy wire for medical guide wire and preparation method of low-cost and high-supporting-performance nickel-titanium shape memory alloy wire

Abstract

The invention provides a low-cost and high-support nickel-titanium shape memory alloy wire for a medical guide wire and a preparation method thereof, the alloy wire is a nickel-titanium binary system, the nickel content is 50.0-52.0 at.%, the austenite phase transformation end temperature is lower than 0 DEG C, the microstructure is a nanocrystalline structure with the average grain size of 30-300 nm, the room temperature unloading platform stress is higher than 200 MPa, and the alloy wire has the advantages that the alloy wire is suitable for being used in the medical guide wire; the preparation method comprises the following steps: carrying out vacuum melting according to the ratio to obtain a cast ingot, sequentially forging, rolling, carrying out multi-pass wire drawing and annealing, and straightening under tension to finally obtain a finished wire. By abandoning third elements such as cobalt and combining component design and nanocrystalline structure regulation and control, the alloy has excellent supporting performance, flexibility and biocompatibility while the material and processing cost is remarkably reduced, and the alloy is particularly suitable for minimally invasive interventional medical guide wires with extremely high requirements for pushing force, passing ability and safety.

Description

Technical Field

[0001] This invention relates to the field of shape memory alloy materials technology, specifically to a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires and its preparation method. Background Technology

[0002] Shape memory alloys, as a class of functional metallic materials with superelasticity and thermo-induced phase change properties, can recover to a pre-set shape after undergoing significant deformation, triggered by temperature or stress. This remarkable mechanical behavior makes them highly promising for applications in many high-tech fields, especially in modern medical devices where the intelligent responsiveness of materials is extremely important. Theoretically, any shape memory alloy with significant superelasticity and good biocompatibility is expected to be used in precision instruments such as vascular interventional guidewires and neural microcatheters to meet the dual requirements of traversing complex paths and adapting to anatomical structures in minimally invasive surgery. However, translating theoretical potential into clinically usable high-performance guidewire materials is far from feasible for all alloys with shape memory effects. The core challenge lies in the fact that medical guidewires, especially ultra-fine guidewires used in coronary artery, cerebrovascular, and peripheral vascular interventions, typically have a diameter of 0.014-0.018 inches. They must achieve a perfect balance of seemingly contradictory mechanical properties on an extremely small spatial scale: excellent "compliance" and sufficient "support". Flexibility ensures that the guidewire tip can conform to the winding and complex branching vascular pathways, allowing for gentle passage without invasiveness or minimal invasiveness, avoiding scratches, dissections, or spasms in the delicate endothelial tissue. Support, on the other hand, provides the guidewire body with sufficient longitudinal stiffness and compressive strength to effectively transmit axial force during advancement, overcoming the resistance and friction of the vessel wall, achieving stable and controllable advancement, and establishing a stable track for subsequent therapeutic instruments. This balance between flexibility and support is a precise control of the material's microscopic deformation mechanism and macroscopic mechanical response. If the material is too flexible, the guidewire is prone to "wrinkling" in the vessel or losing its driving force, failing to reach distal lesions; if it is too rigid, although the pushing force is strong, it easily leads to serious complications such as vascular perforation and intimal damage.

[0003] Currently, to address the aforementioned challenge of balancing flexibility and support, the industry commonly employs nickel-titanium-cobalt alloys with cobalt (Co) as a reinforcing element. This aims to enhance the guidewire's driving force by increasing the plateau stress under the tensile curve. While this method improves mechanical strength to some extent, it also introduces several significant drawbacks. Regarding biosafety, cobalt ions pose additional biosafety risks. If surface integrity is compromised due to processing defects or wear during clinical use, cobalt ions may be released along with nickel ions. Long-term or high-concentration exposure is associated with risks such as myocardial injury, neurotoxicity, and thyroid dysfunction. Although the amount of ions released during guidewire use is typically low, its long-term biosafety boundaries still require further clinical validation, limiting its application in scenarios requiring long-term implantation or repeated use.

[0004] In terms of mechanics and clinical suitability, cobalt-nickel-titanium alloys also face a series of limitations: First, while cobalt increases the alloy's rigidity and pushing force, it significantly reduces its flexibility and bending deformation capacity. In procedures requiring precise bending, such as neurointerventions, this can easily affect the accuracy of the guidewire tip response and increase the risk of vascular intimal damage. Second, cobalt may interfere with the alloy's phase transformation characteristics and phase transformation path, leading to decreased fatigue resistance. Repeated bending can easily cause microcracks or even fracture risks, which may require a second surgery to remove if fractured in vivo. Third, its torsional strength is insufficient. Medical guidewires need to have a certain degree of torsional resistance, meaning that the tip can respond synchronously during rotation to avoid "twisted" deformation. However, the crystal structure and processing technology of nickel-titanium-cobalt alloys (cold-drawn wire forming) result in lower torsional strength than materials such as stainless steel and cobalt-chromium alloys. In complex interventional scenarios requiring rotation and pushing (such as the treatment of chronic coronary occlusion), asynchronous torsion of the guidewire tip can easily occur, affecting surgical efficiency. In addition, in terms of processing and functional integration, the addition of cobalt increases the hardness of the alloy and exacerbates the tendency of work hardening (hardening and embrittlement of the material surface), making diameter control and surface finish more difficult to ensure. If there are tiny scratches on the guidewire surface, they may become the starting point for thrombosis. Furthermore, the phase transformation characteristics of such alloys may lead to dimensional shrinkage during cooling after processing, further reducing dimensional accuracy. Their X-ray imaging performance is weak, far lower than that of precious metals such as tungsten, platinum, and gold. Clinically, they cannot be clearly visualized directly by X-rays, often requiring the welding or embedding of additional imaging markers such as tungsten or platinum-iridium alloys at the guidewire tip. This not only increases process complexity and the risk of interface failure—for example, weak interfacial bonding and micro-gaps at the weld joint can lead to corrosion or breakage—but also the difference in rigidity between the marker and the alloy body can cause abrupt changes in the stiffness of the guidewire tip, increasing the risk of vascular damage. Fourth, some interventional guidewires require coating with antithrombotic drugs such as heparin or anti-inflammatory drugs to reduce postoperative complications. However, the oxide film on the surface of nickel-titanium-cobalt alloys is chemically inert and has poor adhesion to the drug coating, making the coating prone to detachment during operation, resulting in uneven drug release and affecting treatment efficacy. In addition, nickel-titanium-cobalt is a precious metal alloy, and the addition of cobalt further increases the cost of raw materials. At the same time, due to its high processing difficulty, it requires special equipment and process controllers for phase transition temperature and surface quality, resulting in a long production cycle and low yield. Ultimately, its clinical application cost is much higher than that of stainless steel guidewires, thus limiting its application in primary healthcare institutions or low-cost medical scenarios.

[0005] In summary, while existing techniques for reinforcing nickel-titanium guidewires by adding cobalt can improve support, they also introduce significant drawbacks in terms of biosafety, mechanical balance, processing technology, and cost control. In particular, the introduction of cobalt not only increases raw material and manufacturing costs, limiting its widespread use in primary healthcare, but its long-term corrosive behavior and ion release stability in bodily fluids also pose potential risks. Therefore, the medical guidewire field urgently needs a nickel-titanium alloy material and its fabrication process that can achieve high support, high flexibility, good biocompatibility, and lower cost without requiring cobalt or other third elements. This would overcome the multiple limitations of existing technologies and meet the increasingly sophisticated needs of minimally invasive interventional treatments. Summary of the Invention

[0006] The purpose of this invention is to provide a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires and its preparation method. By eliminating third elements such as cobalt and combining composition design and nanocrystalline structure control, the alloy wire can have excellent support, flexibility and biocompatibility while significantly reducing material and processing costs. It is particularly suitable for minimally invasive interventional guidewires with extremely high requirements for pushing force, passability and safety.

[0007] To achieve the above objectives, the present invention proposes the following technical solution:

[0008] A low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires, comprising the following components by atomic mass percentage:

[0009] Ni: 50.0-52.0 at.%, balance Ti;

[0010] The austenite termination temperature of the nickel-titanium shape memory alloy wire is <20℃;

[0011] The nickel-titanium shape memory alloy wire simultaneously meets the following mechanical properties at room temperature:

[0012] The unloading platform stress is 200-270 MPa, and the elongation at break in the drawn state is 3%-30%, while the elongation at break in the annealed state is 20%-35%.

[0013] The microstructure of the nickel-titanium shape memory alloy wire at room temperature is austenite.

[0014] As a preferred embodiment of the present invention, the microstructure of the nickel-titanium shape memory alloy wire is a nanocrystalline structure with an average grain size of 30-300 nm.

[0015] As a preferred embodiment of the present invention, in the nickel-titanium shape memory alloy wire, the area percentage of inclusions is <2%, and the maximum size of a single inclusion is <39μm.

[0016] As a preferred embodiment of the present invention, the diameter of the nickel-titanium shape memory alloy wire is 0.1-0.8 mm.

[0017] This invention also provides a method for preparing a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires, comprising the following steps:

[0018] Step 1: According to the Ni and titanium content ratio, the raw materials are vacuum melted to obtain alloy ingots;

[0019] Step 2: The alloy ingot is forged and rolled sequentially to obtain coarse wire;

[0020] Step 3: Perform multiple drawing processes on the coarse filaments and anneal them under inert gas protection. Control the breaking elongation of the drawn filaments to be 3%-30% and the breaking elongation of the annealed filaments to be 20%-35% to obtain the desired semi-finished filaments.

[0021] Step 4: The semi-finished filament obtained in Step 3 is subjected to tension under inert gas protection to straighten it, thereby obtaining the finished filament.

[0022] As a preferred embodiment of the present invention, the single-pass surface reduction rate of the multi-pass wire drawing process is 8%-50%, and the wire drawing speed is 1-50m / min.

[0023] As a preferred embodiment of the present invention, the annealing process is carried out at a temperature of 350-700℃ and an annealing time of 0.5-30 min.

[0024] As a preferred embodiment of the present invention, during the straightening process, the tension applied to the filament is 10-500 MPa, the straightening temperature is 350-650℃, the straightening speed is 1-30 m / min, and the straightening time is 0.5-10 min.

[0025] As a preferred embodiment of the present invention, the inert gas is at least one of argon and helium.

[0026] As a preferred embodiment of the present invention, in step 1, the obtained alloy ingot comprises, by atomic mass percentage:

[0027] Ni: 50.0-52.0 at.%, balance Ti, and the austenite termination temperature of the alloy ingot is <0℃.

[0028] As can be seen from the above technical solutions, the technical solutions of the present invention provide a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires and its preparation method. Through a series of synergistically optimized technical means, the following comprehensive beneficial effects are achieved:

[0029] First, by eliminating cobalt and other third elements and adopting a pure nickel-titanium binary alloy system, with the nickel content strictly controlled between 50.0% and 52.0%, not only is the use of expensive cobalt metal directly avoided, significantly reducing raw material costs, but the risk of biotoxicity from cobalt ions released during clinical use is also fundamentally eliminated. Simultaneously, this composition design ensures that the austenitic transformation end temperature of the alloy ingot raw material is consistently below 0°C. Subsequent processes such as wire drawing, annealing, and straightening further ensure that the austenitic transformation temperature of the finished alloy wire is below 20°C. This guarantees that the medical guidewire remains in a fully austenitic state within the range of room temperature to body temperature, possessing ready-to-use, stable, and highly elastic properties, providing a reliable and consistent mechanical response basis for surgical procedures.

[0030] Secondly, by precisely controlling the wire drawing and annealing processes—specifically, controlling the wire drawing speed and the reduction rate per pass, and performing annealing at a specific temperature and time under an inert atmosphere—a uniform nanocrystalline structure with a grain size of 30-300 nm was successfully formed inside the finished alloy wire. This microstructural control allows the material to maintain excellent plasticity and flexibility while achieving unloading plateau stresses exceeding 200 MPa. This fundamentally solves the contradiction between support and flexibility in traditional cobalt-added alloys, enabling the guide wire to possess both excellent pushing force and maneuvering capability.

[0031] Furthermore, by optimizing the straightening process and controlling it under specific temperatures and tensions, the nanocrystalline structure of the wire is maintained and further stabilized. This ultimately ensures that the finished alloy wire achieves high unloading platform stress while exhibiting high repeatability and consistency in its phase transformation temperature and mechanical properties. This series of quantifiable and repeatable process controls makes the entire preparation process easy to implement industrially, resulting in a high yield. This reduces both processing difficulty and overall manufacturing costs.

[0032] Furthermore, due to the absence of cobalt, the alloy material exhibits reduced work hardening tendency, making surface quality easier to control. This facilitates obtaining a smooth surface, reducing the risk of thrombosis, and also provides a superior substrate for subsequent drug coating adhesion. Overall, this invention achieves a comprehensive improvement in biocompatibility, mechanical properties, process controllability, and cost-effectiveness while maintaining high performance, providing a more ideal basic material solution for minimally invasive interventional guidewires.

[0033] It should be understood that all combinations of the foregoing concepts and the additional concepts described in more detail below can be considered part of the inventive subject matter of this disclosure, provided that such concepts do not contradict each other. Detailed Implementation

[0034] To make the objectives, technical solutions, and advantages of this invention clearer, the technical solutions of the embodiments of this invention will be clearly and completely described below in conjunction with the embodiments of this invention. Obviously, the described embodiments are only some, not all, of the embodiments of this invention. Based on the described embodiments of this invention, all other embodiments obtained by those skilled in the art without creative effort are within the scope of protection of this invention. Unless otherwise defined, the technical or scientific terms used herein should have the ordinary meaning understood by those skilled in the art to which this invention pertains.

[0035] The terms "first," "second," and similar words used in the specification and claims of this patent application do not indicate any order, quantity, or importance, but are merely used to distinguish different components. Similarly, unless the context clearly indicates otherwise, the singular forms of "an," "a," or "the," etc., do not indicate a quantity limitation, but rather indicate the presence of at least one. Terms such as "comprising" or "including" mean that the element or object preceding "comprising" encompasses the features, integrals, steps, operations, elements, and / or components listed following "comprising" or "including," and do not exclude the presence or addition of one or more other features, integrals, steps, operations, elements, components, and / or collections thereof.

[0036] Firstly, this invention provides a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires. Based on materials science principles, nickel is an element with a strengthening effect. Therefore, to prepare an economical and highly supportive shape memory alloy material, this invention selects a scheme with a higher nickel content, which is essentially a material with a lower phase transition temperature. Considering the required low cost, this invention designs and uses a binary shape memory alloy containing only nickel and titanium, avoiding the use of other metal elements, such as cobalt. While this improves the support of the alloy wire, it introduces other insurmountable defects.

[0037] Specifically, by atomic mass percentage, nickel-titanium shape memory alloy wire comprises the following components: 50.0-52.0 at.% nickel, with the balance being titanium. Due to production limitations, unavoidable impurities may also exist in some specific embodiments. These mainly refer to trace amounts of non-metallic elements, such as O, C, and N, or metallic elements such as Fe and Cr, which are difficult to completely eliminate during raw material smelting and subsequent processing. Their total amount is usually controlled within extremely small quantities and can be ignored.

[0038] Meanwhile, the austenitic termination temperature of the finished nickel-titanium shape memory alloy wire is <20℃, making it suitable for use in the fabrication of interventional medical devices such as medical guidewires. Because medical guidewires are often stored at room temperature (approximately 20-25℃) before surgery, and the time exposed to indoor air during surgery is significant, an austenitic termination temperature <20℃ ensures that the alloy wire is fully in the austenitic phase at room temperature, possessing immediate superelasticity. This eliminates the need for additional heating and activation before operation, and eliminates concerns about performance changes due to indoor temperature fluctuations, improving the convenience and reliability of clinical use. Furthermore, even in a constant-temperature surgical environment, differences in blood flow or different parts of the body can still cause minor local temperature fluctuations. An austenitic termination temperature far below body temperature effectively resists these minor temperature changes, preventing phase structure fluctuations caused by temperature approaching the phase transformation point, and ensuring consistent mechanical properties of the medical guidewire during intravascular advancement, turning, and positioning.

[0039] The microstructure of the nickel-titanium shape memory alloy wire is a nanocrystalline structure with an average grain size of 30-300 nm, and it is an austenitic phase at room temperature. Inclusions in the nickel-titanium shape memory alloy wire are mainly titanium oxides, nitrides, or carbides, originating from atmospheric contamination or raw material residues during the melting process. This invention utilizes vacuum melting and inert gas protection to ensure that the area percentage of inclusions is <2%, and the maximum size of a single inclusion is <39 μm, thereby guaranteeing the consistency of the alloy wire's fatigue performance and processing performance. As a specific embodiment of this invention, the alloy wire diameter can be 0.1-0.8 mm to suit medical guidewires.

[0040] The nickel-titanium shape memory alloy wire simultaneously meets the following mechanical properties at room temperature:

[0041] The unloading platform stress is greater than 200 MPa, preferably 220-270 MPa; and the elongation at break in the drawn state is 3%-30%, while the elongation at break in the annealed state is 20%-35%.

[0042] This invention utilizes the strengthening effect of nickel by appropriately increasing the nickel atomic ratio in the alloy wire. However, the increase in nickel content is not unlimited. As the nickel content in the alloy wire increases, other defects will be introduced, as detailed below.

[0043] Firstly, excessive nickel content leads to an abnormally low phase transformation temperature in the alloy. When the nickel content exceeds 52.0 at.%, the austenitic phase transformation end temperature (Af) of nickel-titanium shape memory alloys drops sharply. For example, at a nickel content of 52.5 at.%, the austenitic phase transformation end temperature may drop below -50°C, while at a nickel content of 55.0 at.%, it drops directly below -100°C. This means that at normal human body temperature (37°C) or typical industrial temperatures, nickel-titanium shape memory alloys will remain in the austenitic phase for extended periods, unable to recover their original shape through heating or unloading, thus directly losing their shape memory effect and superelasticity. The core requirement for nickel-titanium shape memory alloys in the medical field is precisely based on their phase transformation function at room temperature or normal human body temperature; excessive nickel renders the alloy material unusable.

[0044] Secondly, excessive nickel can lead to the precipitation of a brittle second phase in the alloy. If the nickel content increases further, such as exceeding 53.0 at.%, brittle Ni3Ti intermetallic compounds will precipitate during the solidification or heat treatment of nickel-titanium shape memory alloys. This second phase is distributed in needle-like or lamellar shapes at the grain boundaries, which significantly reduces the plasticity and toughness of the alloy, making it prone to cracking during deformation, such as breakage when medical guidewires are bent or rupture when stents are expanded. At the same time, it disrupts the uniformity of phase transformation, causing the shape memory function to be completely lost.

[0045] Third, excessive nickel can exacerbate biosafety risks. Nickel is one of the known allergenic metals, and excessive nickel can directly lead to biosafety hazards. This is the core constraint limiting nickel content in medical settings. The release rate of nickel ions is positively correlated with the nickel content in the alloy material: the higher the nickel content in the alloy material, the more difficult it is to maintain the integrity of its surface oxide film (mainly TiO2, which plays a protective role). This is because when there is an excess of nickel atoms, the oxide film is prone to defects, such as nickel enrichment at grain boundaries leading to uneven oxidation, which makes it more susceptible to corrosion in human body fluids (containing aluminum ions and proteins), resulting in a significantly increased nickel ion release rate. Nickel-titanium shape memory alloys with a nickel content of 52.0-53.0 at.% can control the release of nickel ions below a safe threshold, posing a risk only to a small number of people with allergies. However, if the nickel content exceeds 55.0 at.%, the release of nickel ions may double or even exceed this threshold. This can not only cause local inflammation in most people with allergies, such as endothelial edema and tissue hyperplasia, but may also lead to cytotoxicity due to long-term accumulation, such as affecting myocardial cells and the nervous system. This completely fails to meet the biocompatibility standards for medical implants.

[0046] Fourth, excessive nickel content can lead to an imbalance in the mechanical properties of alloy materials, particularly in the inability to balance strength and plasticity (flexibility). When the nickel content is moderate, such as 50.0-52.0 at.% in nickel-titanium shape memory alloys, it possesses both high yield strength (thanks to the hardness of the austenitic phase) and good plasticity (thanks to the easy deformation of the martensitic phase), meeting the requirements for both rigidity and flexibility in medical guidewire deployment. If the nickel content is slightly excessive, such as at 52.5 at.%, the yield strength of the nickel-titanium shape memory alloy will increase significantly due to the excessive stability of the austenitic phase, but its plasticity will decrease drastically. In medical guidewires, this will manifest as excessive hardness, difficulty in bending, and a tendency to scratch the blood vessel wall. If the nickel content is far excessive, such as >53.0 at.%, the precipitation of the Ni3Ti brittle phase will transform the alloy material from a "plastic material" into a "brittle material," making it prone to direct fracture under external forces, such as breakage during stent expansion, completely rendering it unusable for medical applications. In addition, it will lead to a sharp increase in processing difficulty, a doubling of processing costs, and an extremely low finished product qualification rate, making it uneconomical; it is also detrimental to the control of inclusions, which in turn affects the fatigue performance of alloy materials.

[0047] In the entire process of preparing nickel-titanium shape memory alloy wire, vacuum melting is the only and decisive step that determines the fundamental composition and initial phase transformation characteristics of the material. Subsequent processes such as forging, rolling, drawing, annealing, and straightening mainly serve to control the microstructure of the material, such as grain size and dislocation structure, and the final mechanical properties, such as unloading plateau stress and elongation, through deformation and heat treatment. They do not fundamentally change the nickel-titanium atomic ratio determined by vacuum melting, nor the intrinsic phase transformation temperature of the alloy ingot determined by it, i.e., the austenitic phase transformation end temperature Af.

[0048] Therefore, controlling the nickel content of the alloy ingot material used to prepare nickel-titanium shape memory alloy wire is crucial. The preferred nickel content of the alloy ingot material in this invention is 50.0-52.0 at.%, ensuring that the austenite termination temperature of the alloy ingot material is below 0°C. This is a prerequisite and material basis for obtaining the final finished wire with a fully austenitic structure, stable superelasticity, and excellent comprehensive mechanical properties at room temperature. This invention, by precisely defining the composition and properties of the alloy ingot at the raw material stage, ensures the repeatability and consistency of the final product's performance from the source. This constitutes one of the core features that distinguishes this invention from traditional methods that adjust performance solely through complex post-processing. For example, the size and area of ​​inclusions in the alloy ingot material are key factors affecting the fatigue performance of the finished wire; fewer inclusions in the alloy ingot material mean fewer inclusions in the finished wire. Table 1 below specifically shows the influence of different atomic ratios on the key properties of the alloy ingot.

[0049] Table 1. Atomic proportions and properties of alloy ingots used to prepare finished alloy wires.

[0050]

[0051] In Table 1, the only difference between the examples and Comparative Examples 1 and 2 is the ratio of nickel and titanium atomic masses in the alloy ingots; their processing techniques are identical. Here, Af is the austenitic phase transformation completion temperature of the alloy ingot; inclusion size is the maximum size of a single inclusion in the alloy ingot; inclusion area percentage refers to the area of ​​each of at least five randomly selected fields of view under a metallographic microscope or scanning electron microscope, with an area not less than 0.5 mm². 2 The percentage of the total area of ​​inclusions to the total observed area was calculated, and the average value was taken as the index. Hyperelasticity refers to the ability of the finished alloy wire to undergo significant nonlinear deformation under external force at room temperature (approximately 20-25℃), with strain exceeding 6%, and to essentially recover its original shape after unloading. Its stress-strain curve exhibits distinct loading and unloading plateaus, and the residual strain is less than 0.5%. Finished alloy wires with hyperelasticity demonstrate good bending recovery and pushing stability in medical guidewires. Table 1 shows that minute changes in nickel content significantly affect the austenitic phase transformation end temperature Af of the alloy ingot, as well as the room-temperature hyperelasticity and room-temperature microstructure of the finished alloy wire.

[0052] The atomic ratios of Examples 1 and 2 are both within the limits of this invention. The nickel element is used to strengthen the material. While maximizing the nickel content, the feasibility of smelting is also taken into account. Furthermore, the microstructure of the finished alloy wire at room temperature is austenitic, which has superelasticity.

[0053] Compared to Example 1, Comparative Example 1 had an excessively high nickel content, resulting in an extremely low austenitic phase transformation end temperature for the alloy ingot. The microstructure of the corresponding finished alloy wire at room temperature consisted of matrix + Ni3Ti brittle intermetallic compounds. The precipitation of Ni3Ti brittle intermetallic compounds within the finished alloy wire indicated that the atomic ratio of the alloy ingot material had significantly deviated from the suitable range defined in this invention, leading to an excessively low phase transformation temperature and the precipitation of brittle phases. Furthermore, the resulting finished alloy wire lacked superelasticity. In addition, the inclusions in the alloy ingot were very large, which was also related to the precipitation of brittle intermetallic compounds.

[0054] Compared to Example 1, Comparative Example 2 had a significantly lower nickel content, causing the austenitic transformation end temperature of the alloy ingot to rise to 110°C. At this temperature, the finished alloy wire exhibited a martensitic microstructure at room temperature, meaning it could not recover its original shape after deformation, indicating a loss of superelasticity. Only by heating it to above the martensitic transformation initiation temperature As could the alloy wire recover its original shape. Furthermore, the inclusion size of the alloy ingot was larger than in Examples 1 and 2. This is because Comparative Example 2, being rich in titanium, was more prone to oxidation / micro-area component segregation during the smelting process, resulting in slightly larger inclusion size and area percentage compared to the examples.

[0055] On the other hand, the present invention also provides a method for preparing a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires, which mainly includes processes such as vacuum melting, forging, rolling, drawing, annealing, and straightening.

[0056] As mentioned earlier, in the entire process of preparing nickel-titanium shape memory alloy wire, except for vacuum melting, the other processes basically do not affect the atomic ratio of the material / the atomic ratio of the finished product. Therefore, it is only necessary to design a reasonable nickel-titanium atomic ratio in the vacuum melting stage, such as using the alloy ingot with a nickel content of 50.0-52.0 at.% and the balance of Ti as the raw material for preparing the finished alloy wire.

[0057] Unlike existing technologies that improve material support by adding other metallic elements, this invention achieves this through alloy atomic ratio design, controlled wire drawing, annealing, and controlled straightening processes. The main preparation steps are as follows:

[0058] Step 1: According to the Ni and titanium content ratio, the raw materials are vacuum melted to obtain an alloy ingot, wherein the nickel content of the alloy ingot is 50.0-52.0 at.%, and the balance is Ti, the area percentage of the inclusions is <2%, and the maximum size of a single inclusion is <39μm.

[0059] Step 2: The alloy ingot is forged and rolled sequentially to obtain coarse wire.

[0060] Step 3: Perform multiple drawing processes on the coarse filaments and conduct intermediate annealing at a specific temperature under inert gas protection to control the breaking elongation of the drawn filaments and the breaking elongation of the annealed filaments, thereby obtaining the desired semi-finished filaments.

[0061] The drawing and annealing parameters in this step significantly affect the performance of nickel-titanium shape memory alloy wires and are crucial for their suitability for medical guidewires. To prepare a qualified finished alloy wire, it is necessary to control the drawing-annealing process parameters to obtain a finished alloy wire with a nanocrystalline microstructure. The grain size of the nanocrystals is 30-300 nm. This nanocrystalline structure provides the microscopic guarantee for achieving both high support and flexibility in medical guidewires. A grain size smaller than 30 nm may lead to processing difficulties and reduced elongation; a grain size larger than 300 nm results in insufficient strengthening effect, making it difficult to achieve an unloading plateau stress of over 200 MPa. Simultaneously, the final drawn wire should have a breaking elongation of 3%-30%, preferably 4%-25%; and the annealed wire should have a breaking elongation of 20%-35%.

[0062] Specifically, the preferred wire drawing process parameters are as follows: multi-pass wire drawing, with a single pass surface reduction rate of 8%-50% and a wire drawing speed of 1-50m / min.

[0063] Optionally, lubricating oil can be used for lubrication and cooling during the wire drawing process. Specifically, oil-based or water-based lubricants can be used. During use, the die holder and mold are completely immersed in the lubricant to ensure that the heat generated during the wire drawing process can be carried away in time. The lubricant circulation system is equipped with a filter screen with a pore size of <1μm to ensure that the lubricant is free of large impurity particles.

[0064] Specifically, the annealing process parameters are as follows: annealing temperature is 350-700℃; annealing time is 0.5-30 min; annealing atmosphere: argon, helium, or a mixture of both, or other inert gases well known to those skilled in the art. If the annealing parameters are selected appropriately, the filament will show no significant color change after one annealing process following drawing.

[0065] The mechanical properties of the alloy wires during the drawing and annealing processes were tested at room temperature using existing mechanical tensile testing equipment in the field, and the wire elongation and breaking strength were recorded.

[0066] Step 4: The semi-finished wire obtained in Step 3 is straightened under inert gas protection to obtain the final finished wire. The straightening process parameters determine the final properties of the finished alloy wire.

[0067] Specifically, the preferred straightening process parameters are as follows: straightening temperature is 350-650℃, preferably 450-600℃; the tension applied to the wire is 10-500MPa; the straightening speed is 1-30m / min; the straightening time is 0.5-10min; and the cooling method is water cooling. The final straightened alloy wire exhibits an austenitic phase transformation end temperature <20℃ and an unloading platform stress >200MPa.

[0068] Table 2. Process parameters and properties of nickel-titanium shape memory alloy wire drawing and annealing

[0069]

[0070] In Table 2, Examples 2-1 and 2-2 represent two performance states of Example 2 in Table 1 after treatment with different drawing and annealing parameters. Comparative Examples (2-4) to (2-7) also represent the performance states of Example 2 in Table 1 after treatment with different drawing and annealing parameters. Compared to Examples 2-1 and Comparative Example 2-3, Comparative Example 2-4 aims to illustrate the effect of the choice of drawing rate on whether the wire breaks, Comparative Example 2-5 aims to illustrate the effect of drawing rate on elongation and grain size, Comparative Example 2-6 aims to illustrate the effect of the choice of annealing temperature on the color of the wire, and Comparative Example 2-7 aims to illustrate the effect of the choice of annealing time on elongation and grain size. The first fracture elongation is the fracture elongation in the drawn state, and the second fracture elongation is the fracture elongation in the annealed state; " / " indicates that there is no corresponding operation and related parameters for this item. In the single-pass processing of alloy wire, drawing and annealing operations are included, but they are not performed simultaneously, but sequentially. For example, in Example 2-1, when the alloy wire is being drawn, there is no annealing operation at the same time. Therefore, the data for annealing temperature, annealing time, and second fracture elongation in the same row are all absent, indicated by " / ". Annealing is only performed after the drawing process is completed. Therefore, when parameters such as annealing temperature and annealing time appear, it indicates that annealing is being performed. At this time, there are no parameters in the drawn state. Therefore, there are no parameters such as drawing speed and reduction rate in the corresponding row, indicated by " / ". The other Examples 2-2 and Comparative Examples (2-3) to (2-7) are similar.

[0071] As shown in Table 2, compared to Example 2-1, Comparative Example 2-3 had the same reduction rate during the wire drawing process, but its wire drawing speed was too fast, resulting in a certain degree of reduction in grain size after wire drawing, and a slight decrease in both the first fracture elongation and the second elongation. More seriously, Comparative Example 2-3 experienced wire breakage during the wire drawing process. This was because the excessively fast wire drawing speed prevented the heat generated during wire drawing from being sufficiently carried away by the coolant, causing micro-oxidation and resulting in a bluish-purple appearance in some areas. The annealing process parameters of Comparative Example 2-3 were the same as those of Example 2-1, but wire breakage also occurred. This may be due to the micro-cracks that had already formed inside the alloy wire during the previous wire drawing process.

[0072] Compared to Example 2-1, Comparative Examples 2-4 used the same drawing speed, but the reduction ratio increased to 60%. This excessive reduction ratio resulted in smaller grain size after drawing, and a decrease in both the first fracture elongation and the second elongation. The drawing process not only resulted in severe wire breakage but also extremely high wire material loss. The annealing process of Comparative Examples 2-4 was the same as that of Example 2-1, but wire breakage also occurred. This is likely due to microcracks forming inside the alloy wire during the preceding drawing process.

[0073] Compared to Example 2-1, Comparative Examples 2-5 had the same wire drawing speed, but the reduction rate was too small. This resulted in insufficient grain refinement of the alloy wire during the wire drawing process, leading to a significantly larger grain size than that of Example 2-1, which also resulted in low processing efficiency. Although the annealing process parameters were the same as those of Example 2-1, the failure of the preceding wire drawing process to process the grain size to a suitable range meant that this pass only changed the diameter of the alloy wire and did not improve its performance, nor did it prepare the necessary microstructure size for the subsequent straightening process.

[0074] Compared to Example 2-1, Comparative Documents 2-6 have the same wire drawing process parameters, but the annealing temperature is too high, resulting in severe oxidation of the alloy wire and the appearance of a blackish-gray oxide layer on its surface. Although the second fracture elongation is improved, the grain size is too large.

[0075] Compared to Example 2-1, Comparative document 2-7 differs in that the annealing time is too long, resulting in severe oxidation of the alloy wire, with a dark blue oxide layer appearing on its surface and excessive grain growth.

[0076] Table 3 Straightening process parameters and properties of nickel-titanium shape memory alloy wire

[0077]

[0078] In Table 3, the phase transformation temperature Af is the austenitic phase transformation end temperature Af of the finished alloy wire.

[0079] As can be seen from Table 3, Comparative Example 2-2-4 has the same nickel-titanium atomic ratio as Example 2-2-1, and the same wire drawing and annealing process parameters. The only difference is the straightening temperature and straightening speed. Comparative Example 2-2-4 has a higher straightening temperature and a slower straightening speed, which significantly reduces the stress on the unloading platform. This is because the straightening temperature is too high and the straightening time is too long, which leads to the rapid elimination of the deformed structure of the alloy wire during the straightening process.

[0080] Compared with Example 2-2-1, Comparative Example 2-2-5 has the same nickel-titanium atomic ratio and the same wire drawing and annealing process parameters. The only difference is that the wire tension is smaller than that of Example 2-2-1, which ultimately leads to a decrease in the stress of the unloading platform and a slight decrease in the phase transformation temperature Af.

[0081] While the present invention has been disclosed above with reference to preferred embodiments, it is not intended to limit the invention. Those skilled in the art can make various modifications and refinements without departing from the spirit and scope of the invention. Therefore, the scope of protection of the present invention shall be determined by the claims.

Claims

1. A low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires, characterized in that, Based on atomic mass percentage, it includes the following components: Ni: 50.0-52.0 at.%, balance Ti; The austenite termination temperature of the nickel-titanium shape memory alloy wire is <20℃; The nickel-titanium shape memory alloy wire simultaneously meets the following mechanical properties at room temperature: The unloading platform stress is 200-270 MPa, and the elongation at break in the drawn state is 3%-30%, while the elongation at break in the annealed state is 20%-35%. The microstructure of the nickel-titanium shape memory alloy wire at room temperature is austenite.

2. The low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires according to claim 1, characterized in that, The microstructure of the nickel-titanium shape memory alloy wire is a nanocrystalline structure with an average grain size of 30-300 nm.

3. The low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires according to claim 1, characterized in that, In the nickel-titanium shape memory alloy wire, the area percentage of inclusions is <2%, and the maximum size of a single inclusion is <39μm.

4. The low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires according to claim 1, characterized in that, The diameter of the nickel-titanium shape memory alloy wire is 0.1-0.8 mm.

5. A method for preparing a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires as described in any one of claims 1-4, characterized in that, Includes the following steps: Step 1: According to the Ni and titanium content ratio, the raw materials are vacuum melted to obtain alloy ingots; Step 2: The alloy ingot is forged and rolled sequentially to obtain coarse wire; Step 3: Perform multiple drawing processes on the coarse filaments and anneal them under inert gas protection. Control the breaking elongation of the drawn filaments to be 3%-30% and the breaking elongation of the annealed filaments to be 20%-35% to obtain the desired semi-finished filaments. Step 4: The semi-finished filament obtained in Step 3 is subjected to tension under inert gas protection to straighten it, thereby obtaining the finished filament.

6. The method for preparing a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires according to claim 5, characterized in that, The single-pass reduction rate of the multi-pass drawing process is 8%-50%, and the drawing speed is 1-50m / min.

7. The method for preparing a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires according to claim 6, characterized in that, During the annealing process, the annealing temperature is 350-700℃ and the annealing time is 0.5-30min.

8. The method for preparing a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires according to claim 7, characterized in that, During the straightening process, the tension applied to the filament is 10-500 MPa, the straightening temperature is 350-650℃, the straightening speed is 1-30 m / min, and the straightening time is 0.5-10 min.

9. The method for preparing a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires according to claim 5, characterized in that, The inert gas is at least one of argon and helium.

10. The method for preparing a low-cost, high-support nickel-titanium shape memory alloy wire for medical guidewires according to claim 5, characterized in that, In step 1, the obtained alloy ingot comprises, by atomic mass percentage: Ni: 50.0-52.0 at.%, balance Ti, and the austenite termination temperature of the alloy ingot is <0℃.

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