WO2004022118A1 - Biomaterial member - Google Patents

Abstract

It is intended to provide a biomaterial member made of an amorphous alloy which shows a polarization resistance in degassed Hank’s solution of 4x106 Ωcm2 or more and an immersion potential-pitting corrosion potential window of 0.25 V or more. In an appropriate embodiment, the above biomaterial is made of an amorphous alloy having a Young’s modulus of 100 GPa or less. Such a biomaterial is free from Ni serving as an allergy source and highly resistant to corrosion due to biological fluids. In addition, it has a high strength and a low modulus of elasticity, scarcely reacts with a living body and can be easily drawn out after the completion of a treatment. In the case of using as a bone marrow nail or the like, it scarcely causes problems such as so-called stress shield and pulmonary infarction. Owing to these characteristics, it is the most desirable artificial bone material to be used in, for example, a bone marrow nail (1) which is embedded in a bone A and then fixed with an interlocking nail (a horizontally fastening screw) (2), a bone connection plate, or a bone fastening screw.

Description

 Description Biomaterial components Technical field

 The present invention relates to a biomaterial made of an amorphous alloy (metallic glass), and more particularly, to an orthopedic spinal fixation material such as an intramedullary nail or an osteosynthesis plate, a fracture fixation material, an artificial joint, or an intervertebral joint. The present invention relates to a biological material member useful for dental crowns, inlays, crowns, artificial beds, artificial dental roots, orthodontic wires, etc., and surgical instruments for general surgery. Background art

 The requirements for medical materials include non-toxicity, non-carcinogenicity, non-allergenicity, suitability for living tissues, and biocompatibility for safety, as well as medical functionality for enhancing medical effects. And the like. Medical functionality includes mechanical properties (mechanical strength, fatigue resistance, wear resistance) and chemical properties (corrosion resistance, biofluid corrosion resistance).

 In recent years, with the progress of medicine, metal materials such as various artificial bones and artificial organs have been often left in vivo for a long time. Along with that, various medical troubles caused by such materials have occurred.

 For example, as a material for forming a hard tissue replacement device such as an artificial hip joint or an artificial tooth root, a Co_Cr-based alloy, stainless steel, or a titanium alloy has been conventionally used, and various titanium alloys have been proposed. (See, for example, JP-A-58-124244).

However, stainless steel SUS316L and Ni-Ti alloys contain Ni and are a source of allergy. In fact, recent reports have shown that titanium alloys have also caused allergies, and the use of stainless steel or titanium alloys is not always optimal for living organisms. In addition, titanium alloy is used for intramedullary nails, etc. In addition to the problem that it is difficult to remove the bone after it is fused with the bone, it is difficult for stress to be transmitted to the bone tissue, and the load on the bone is cut off, so that bone resorption is promoted and bone atrophy is used. The so-called stress shield problem (problem of load cut-off), which causes the problem, is likely to occur.

 On the other hand, various proposals have been made for the use of organic polymer materials as biomaterials, for example, using biodegradable and absorbable borylactide and its copolymers, and copolymers with polyglycolic acid and the like. An osteosynthesis material (see Japanese Patent Application Laid-Open No. 3-176660) or a osteosynthesis material composed of a composite material containing calcium phosphate glass fiber and an organic polymer material that does not inhibit biocompatibility. Kaihei 5 — 1 46502) has been proposed.

 However, in addition to the problem of corrosion resistance to body fluids, organic polymer materials have a fatal drawback of low strength. For this reason, the size of the bone must be very large.For example, when a thick intramedullary nail is inserted into the bone marrow, the fat that overflows from the bone marrow mixes into the blood, thereby forming a thrombus in the lungs. There is a problem that it is easy to cause infarction. It can also occupy the medullary cavity extensively, block blood flow in the bone marrow, and delay bone union.

 The present invention has been made in view of the above-mentioned circumstances, and has as its object the purpose of not including Ni as an allergic source, having excellent corrosion resistance to biological fluids, high strength, and low elastic modulus. Another object of the present invention is to provide a biomaterial member that is difficult to react with a living body, is easy to remove after treatment, and hardly causes problems such as stress shield and pulmonary infarction when used as an intramedullary nail or the like. Disclosure of the invention

In order to achieve the above object, according to the present invention, a polarization resistance in a degassed Hanks solution is 4 × 10 ⁶ Ωcm ² or more, and an immersion potential-pitting potential window is 0.25 V or more. A biomaterial member comprising an amorphous alloy of the above is provided. In a preferred embodiment, the biomaterial is made of an amorphous alloy having a Young's modulus of 10 OGPa or less. The biomaterial member of the present invention does not contain Ni as an allergic source, has excellent corrosion resistance to biological fluids, has high strength and a low elastic modulus, is difficult to react with a living body, and is removed after the treatment. When used as an intramedullary nail or the like, problems such as so-called stress shielding and pulmonary infarction are unlikely to occur.

The biomaterial member may be an amorphous alloy having a polarization resistance in a degassed Hanks solution of 4 × 10 ⁶ Ωcm ² or more, and an immersion potential-pitting potential window of 0.25 V or more. Since it is preferably made of an amorphous alloy having a Young's modulus of 100 GPa or less, it has excellent in vivo corrosion resistance, excellent specific strength, low Young's modulus, and physical properties closer to bone. . For this reason, a smaller osteosynthesis material with the same strength can be manufactured, and it is less susceptible to stress shields, which is advantageous for bone fusion, and a conventional thick intramedullary nail is inserted into the bone marrow. In this case, the fat that overflows from the bone marrow mixes into the blood, thereby forming a blood clot in the lungs, which makes it easier to cause pulmonary infarction. The effect of reducing the increase in intramedullary pressure can be obtained. In addition, the space occupation in the bone marrow can be minimized, blood vessels in the bone marrow can be preserved to a maximum extent, and bone formation in the bone marrow is hardly inhibited.

 The most preferable application mode of the biomaterial member of the present invention is at least one of an intramedullary nail for orthopedic surgery, a osteosynthesis plate, a locking nail, and a bone fixing screw. BRIEF DESCRIPTION OF THE FIGURES

Figure 1 is a schematic diagram of order to explain the method for obtaining the polarization resistance R _p from the polarization curve of the corrosion potential neighbors.

2, Z r _5. Is a graph showing a p H-dependent immersion potential of A ₁ 1 _Q C u _{4 Q} amorphous alloy.

 FIG. 3 is a schematic partial cross-sectional front view showing an embodiment in which the biomaterial made of the amorphous alloy of the present invention is applied to an intramedullary nail.

FIG. 4 shows the application of the biomaterial made of the amorphous alloy of the present invention to an osteosynthesis plate. It is a schematic front view showing an embodiment.

_{5, Z r 5 0 A 1 1} 0 C u 4 0 § Roh one de polarization curve degassed H Anks solution of pure T i and SUS 3 1 6 L steel is comparison with amorphous alloys This is a graph showing. BEST MODE FOR CARRYING OUT THE INVENTION

As described above, biomaterials used as intramedullary nails, osteosynthesis plates, etc. are not only non-toxic, non-allergic, but also suitable for blood, such as biocompatibility for safety, as well as mechanical properties ( Mechanical strength, fatigue resistance, and abrasion resistance) and chemical properties (corrosion resistance, corrosion resistance to biological fluids) are required. polarization resistance is 4 x 1 0 ⁶ Ω cm ² or more, immersion potential - from the pitting potential window 0 2 5 V or more amorphous alloy, preferably a Young's modulus of 1 0 0 GP a less amorphous alloy. When fabricated, they have almost satisfied the above-mentioned required properties and found that they are most suitable as a biomaterial, and have completed the present invention.

That is, the first feature of the biomaterial member of the present invention is that it is an amorphous alloy having a polarization resistance of 4 × 10 ⁶ Ωcm ² or more in a degassed Hanks solution.

 If a metal material corrodes in a living body, metal ions are eluted and a corrosion product is formed. The metallic material does not show toxicity such as allergy or carcinogenesis as it is, but in order to exhibit toxicity, it is eluted as metal ions by corrosion and binds to biomolecules in the form of ions or metal salts, or is similar to abrasion powder. It must be in a proper form. In addition, the destruction of metallic materials in vivo is attributed to fatigue and fretting fatigue.However, these are not singly occurring, but are phenomena related to corrosion such as corrosion fatigue and fretting corrosion. . As described above, the corrosion phenomena of metallic materials in the living environment is important from the viewpoint of the toxicity and durability of the material, and measurement for clarifying the corrosion phenomena in the living environment in vitro is required.

The choice of a solution for use in measuring biomaterial corrosion is extremely important. Ideally, the test should be performed under the same conditions as the use environment, but the composition and the exact same as the body fluid It is virtually impossible to use the solution in a state in vitro. Since the corrosion of metallic materials generally involves C1—, the simplest solution for measuring corrosion is the 0.9% NaCl solution (physiological saline). However, this solution does not contain many ions contained in body fluids. In addition, Ringer's solution does not contain phosphate ions, whereas phosphate buffer solution (PBS (-)) does not contain cations such as Ca ²⁺ and Mg ²⁺ . In contrast, Hanks' solution has a composition close to that of the extracellular solution, as shown in Table 1 below.

そこで、 本発明では、 試験液として脱気した H a nk s液を用い、 この中 での分極抵抗が 4 X 10⁶ Ω cm²以上の非晶質合金を用いるものである。 周知のように、 腐食反応の評価に用いられる分極抵抗: _pは、 微小な電圧 変分 Δ ?7又は電流変分 Δ iを与えたときに得られる電流又は電圧変化との比 (厶？ ?/Δ ;ί) であり、 図 1に示すように分極曲線の腐食電位 E _c。_r付近の 接線の勾配から R_pを求める方法、 I △?? 1 < 1 OmVの定電位ステップを 与えて電流を読む方法、 あるいは I Δτ? Iく 1 OmVになるような電流を選 び、定電流ステップを与え、電位変化を読む方法などが広く行なわれており、 いずれも可能であるが、 簡便で精度良く測定できることから図 1に示すよう に分極曲線の = 0付近の接線の勾配から R_pを求める方法が好ましい。 本発明では、 脱気した H a nk s液中での分極抵抗が 4 x 10⁶ Ω c m²以 上の非晶質合金を用いるが、分極抵抗 R _pが大きくなるほど腐食電流密度 I。

Therefore, in the present invention, a degassed Hanks solution is used as a test solution, and an amorphous alloy having a polarization resistance of 4 × 10 ⁶ Ωcm ² or more is used. As is well known, the polarization resistance used for evaluating the corrosion reaction: _p is the ratio of the current or voltage change obtained when a small voltage variation Δ? 7 or a current variation Δi is given. / Δ; ί), and the corrosion potential E _c of the polarization curve as shown in FIG. Select the method of obtaining R _p from the slope of the tangent near _r , the method of reading the current by giving a constant potential step of I △? 1 <1 OmV, or the current that gives I Δτ? There are widely used methods such as applying a constant current step and reading the potential change.Either method is possible.However, since the measurement can be performed simply and accurately, as shown in Fig. 1, the slope of the tangent line near = 0 in the polarization curve A method for determining R _p is preferred. In the present invention, the polarization resistance in the degassed Hanks solution is 4 × 10 ⁶ Ωcm ^{2 or} less. Although the above amorphous alloy is used, the corrosion current density I increases as the polarization resistance _Rp increases.

. _r decreases and the corrosion rate decreases. Therefore, there is no point in setting the upper limit of the polarization resistance, and it is only necessary to satisfy the condition that the polarization resistance is 4 × 10 ⁶ Ωcm ² or more.

 Next, a second feature of the biomaterial member of the present invention is that the biomaterial member is an amorphous alloy having an immersion potential-pitting corrosion potential window of 0.25 V or more.

 Here, the immersion potential-pitting potential window is simply the potential difference between the pitting potential and the immersion potential on the anodic polarization curve, and indicates the allowable range of an environmental change that causes pitting.

 The immersion potential is the potential of a metal when the metal is immersed in a solution and no action is applied from the outside. Here, if there is an environmental change such as pH change of the solution, the immersion potential of the metal changes, and if the immersion potential after the environmental change becomes higher than the pitting potential, pitting occurs. Therefore, the larger the difference between the pitting potential and the immersion potential, the wider the allowable range for metal potential changes due to environmental changes. As described above, since an environmental change with respect to a metal can be simulated as a change in the potential of the metal, in the present invention, the pitting corrosion resistance of the metal with respect to the environmental change is defined by an immersion potential-porosity potential window by a polarization test. Things.

 In the present invention, the immersion potential-pitting potential window is specified to be 0.25 V or more for the following reason.

Single-roll liquid in a phosphoric acid-cunic acid solution with different pH (chloride ion concentration: 0 mol / L, pH: 2.2-8.0, dissolved oxygen: degassed, temperature: 310 K) ribbon-like Z r _5, which was produced by quenching method. A 1 i. C u _4. As shown in FIG. 2, when the immersion potential of the amorphous alloy having the composition shown in FIG. 2 was changed from pH 7.5 (normal) to pH 2.2, the immersion potential of the alloy was about 0.2. 2 5 V higher. Since the pH of the inflamed site is known to be about 5.2, it can be said that in this amorphous alloy, the change in immersion potential due to inflammation or the like is less than 0.25 V. Therefore, if the immersion potential-pitting potential window is 0.25 V or more, preferably 0.3 V or more, it is determined that pitting does not occur even by environmental changes in the living body. In the present invention, an amorphous alloy having an immersion potential-pitting potential window of 0.25 V or more is used. Since it is determined that pitting does not occur due to environmental changes, there is no point in setting the upper limit of the immersion potential-pitting potential window, and only the condition of 0.25 V or more has to be satisfied.

 As described above, the biomaterial for orthopedic surgery of the present invention is made of an amorphous alloy satisfying a specific corrosion resistance condition. It also has the advantage of the above. The characteristics of an amorphous alloy (metallic glass) vary depending on the combination of its constituent elements.

① Higher corrosion resistance than crystalline metal of the same composition,

 (2) The Young's modulus / strength ratio is small, making it more useful for biomaterial applications.

③ Can be formed from a supercooled liquid zone that is much lower than the melting temperature

 The structure has super fine transferability,

The advantages based on these characteristics can be obtained.

 Among the above-mentioned properties, the Young's modulus is important in the properties required for biomaterials for orthopedic surgery, especially for intramedullary nails, osteosynthesis materials, interlocking nails, and screws for bone fixation. Below, preferably 90 GPa or less is desirable. However, it is desirable to have a Young's modulus of about 20 GPa or more, which is close to that of bone.

 The Young's modulus at the densest bone part in the living body is at most about 20 GPa, which is about 1/10 that of stainless steel. For example, in the case of an osteosynthesis member, if the Young's modulus is high, the load is not transmitted to the bone but is applied to the metal plate, and stress is not transmitted to the bone tissue (so-called stress shield problem). As a result, bone union may be delayed, and the fracture may be weakened even if it is union. Even harder artificial materials can compress the bone and cause impaired blood flow at the interface between the member and the bone.

According to the study of the present inventors, it was found that the smaller the Young's modulus of the intramedullary nail and the osteosynthesis material, the better the bone fusion. As shown in the test examples described below, When implanted in a rat femur as an intramedullary nail for fixing a fracture, stainless steel (SUS316L) with a Young's modulus of about 200 GPa or Ti-6A with a lower Young's modulus than that Amorphous alloys had the greatest amount of bone formation compared to 1-4 V alloys. This is because if the material has a low modulus of elasticity, the bone will share the load appropriately and this will promote bone formation. In addition, calcium phosphate was observed on the surface of the titanium alloy after extraction and sulfide was observed on stainless steel, whereas no product was observed on the amorphous alloy. When calcium phosphate is formed, as in titanium alloys, it strengthens the bond between bone and the material, making it difficult to remove after bone fusion.However, in the case of amorphous alloys, such products are not generated. However, it becomes easy to remove after bone fusion, and it is most suitable as an intramedullary nail or osteosynthesis material.

 For reference, Young's modulus and tensile strength are shown in Table 2 below for some examples of amorphous alloys.

 Table 2

前述したように、 従来の金属系生体用骨接合材料は生体適合性、 生体安全 性、 耐食性、 耐久性などに問題があることが指摘されており、 また、 骨に比 較してヤング率が大きいため、 骨接合材料として使用した場合、 骨形成に必 要な荷重が遮断され、 時に骨癒合が遷延する傾向があった。

As mentioned above, it has been pointed out that conventional metal-based osteosynthesis materials for living bodies have problems in biocompatibility, biosafety, corrosion resistance, durability, etc., and have a Young's modulus that is lower than that of bone. Due to its large size, when used as an osteosynthesis material, the load required for bone formation was interrupted, and bone fusion sometimes prolonged.

In contrast, amorphous alloys have excellent specific strength, low Young's modulus, and have physical properties closer to bone. For this reason, it is possible to produce an osteosynthesis material having the same strength and a smaller size (for example, the diameter of a hole having the same strength as a titanium alloy rod having a diameter of about 20 mm is about 15 to 16 mm) It is considered that it is hardly affected by load shedding and is advantageous for bone union. Therefore, when a thick intramedullary nail is inserted into the bone marrow as in the past, fat overflowing from the bone marrow enters the blood. In addition, there is less of a problem that blood clots are formed in the lungs and pulmonary infarction is likely to occur, and the size can be reduced, so that the increase in intramedullary pressure during surgery can be reduced. In addition, the space occupation in the bone marrow can be minimized, blood vessels in the bone marrow can be preserved to a maximum extent, and bone formation in the bone marrow is hardly inhibited.

 In this way, even in the case of a material composed of the same metal element, the amorphous state without grain boundaries is often higher in corrosion resistance than the crystalline metal, and is extremely advantageous as a material for medical use, particularly as an osteosynthesis material. is there.

 As described above, amorphous alloys have lower Young's modulus and higher strength than crystalline metals, but these are very valuable properties for biomaterials. The biomaterial member of the present invention is most suitable as an artificial bone material because of its low Young's modulus. However, for other uses, for example, for shaping a spinal column fixing material, a fracture fixing material, an artificial joint, an intervertebral spacer, etc. It is useful as a material for various medical uses, such as surgical materials, dental materials such as crowns, inlays, crowns, prostheses, artificial roots, orthodontic wires, and general surgical materials such as surgical instruments.

The biomaterial member of the present invention has a polarization resistance in the degassed Hanks solution of 4 × 10 ⁶ Ωcm ² or more, and an immersion potential-pitting potential window of 0.25 V or more, preferably In addition, any substantially amorphous alloy material containing an amorphous phase having a Young's modulus of 100 GPa or less and at least a volume fraction of 50% or more can be applied without particular limitation. The present invention can be suitably applied to an amorphous alloy having a composition represented by any one of the formulas (1) and (2). If the volume ratio is less than 50%, the invention can be applied to an amorphous alloy containing nanocrystals or quasicrystals.

General formula (1):

(However, M ¹ is one, two or three elements selected from Zr, Hf and Ti; M ² is Cu, Fe, Co, Mn, Nb, V, Cr, Zn, At least one element selected from the group consisting of Al, Sn and Ga; M ³ is at least one element selected from the group consisting of B, C, N, P, Si and 0; and M ⁴ is Ta , W And at least one element selected from the group consisting of Mo, M ⁵ is Au, P t, at least one element selected from the group consisting of Pd and Ag, a, b, c, d, and e it It is in atomic%, 25≤a≤85, 15≤b≤75, 0≤c30, 0≤d≤15, 0≤e≤15. )

 General formula (2): A 1: l ~ 10%, Ga: 0.5 ~ 4%, P: 9 ~ 15%, C: 5 ~ 7%, B: 2 ~ 10%, S in atomic% i: 0 to: L 5%, Ge: 0 to 4%, Fe: Fe-based amorphous alloy having the balance of the composition.

In the amorphous alloy represented by the general formula ^(1), M 1 is an essential-based metal element order to form an amorphous, M ² decreases the melting point by combining the M ^1, eutectic This has the effect of easily causing supercooling during solidification, and can facilitate the formation of amorphous. M ³ represents improves the corrosion resistance of the amorphous metal, supercooled liquid region (glass transition region) ΔΤχ (= Τχ- T g: Τχ crystallization temperature, Tg is the glass transition temperature) broadening the temperature range of effects And can be stabilized against crystallization. However, no longer form an amorphous and M ³ element is too many. M ⁴ promotes passivation of the surface of the amorphous phase and improves corrosion resistance. M ⁵ is improves the corrosion resistance, is suppressed oxidation of the molten metal by the flux effect, the degree of supercooling for reducing heteronuclear production formation site increases, improving the amorphous formation capability. Among these, an amorphous alloy having a glass transition region having a temperature width of 30 K or more is preferable.

 Hereinafter, some of the embodiments of the present invention will be described with reference to the accompanying drawings, and test examples for specifically confirming the effects of the present invention will be shown. However, it goes without saying that the present invention is not limited to the following embodiments and test examples.

 FIG. 3 shows an embodiment in which the biomaterial made of the amorphous alloy of the present invention is applied to an intramedullary nail. After being implanted in bone A, the intramedullary nail 1 is fixed with an interlocking nail (side stop screw) 2. Here, the intramedullary nail 1 and the inner rocking nail 2 are made from the amorphous alloy of the present invention.

Next, Fig. 4 shows the application of a biomaterial made of the amorphous alloy of the present invention to an osteosynthesis plate. 2 shows an embodiment according to the invention. The joint plate 3 is applied to the bone A fractured at the fracture site X, and is fixed by the osteosynthesis plate fixing screw (screw) 4. Here, the joint plate 3 and the screw 4 for fixing the osteosynthesis plate are made of the amorphous alloy of the present invention.

 Test example 1

Degassed _{Z r 50 A 1 x 0 C} u 40 amorphous alloy in H Anks solution was subjected to anode polarization test stainless steel SUS 3 1 6 L and pure T i. The test conditions are as follows.

 5ϊ \ ι fee:

(1) Z r _5. Al ₁₀ 〇 11 ₄ . Amorphous alloy ribbon, 3 types (η = 3) The surface on the argon gas side during quenching in the quenched state was used for measurement.

 (2) Stainless steel SUS316L (one commercial product (η = 1))

 The surface polished with # 600 SiC abrasive paper was used for measurement.

 (3) Pure Ti (commercially available, two types of JIS (n = 2))

 The surface polished with # 600 SiC abrasive paper was used for measurement.

 Figure 5 shows the measured polarization curves. Table 3 shows the corrosion characteristic values obtained from the polarization curves.

 Table 3

図 5に示される分極曲線から明らかなように、 Z r^Al ^ C iiA。非晶 質合金は、 脱気した H a nk s液中でのアノード分極に伴い速やかに自己不 働態化し、 一定の不働態維持電流を示した後、 孔食を発生した。 通常の体液 環境での各合金の電位は、 本分極曲線における自然浸漬電位（E。_{p en})付近 であると考えられる。そこで、 E。_{p e n}での材料の耐腐食性に相応する分極抵 抗 （R_p) を、 E。_{p e n}での電流密度の傾きから求めた。 本非晶質合金の R_p は、 SUS 3 16 L鋼よりも一桁高く、純 T iよりも数倍高かったことから、 本非晶質合金の耐食性は S US 3 1 6 L鋼よりも非常に高く、 純 iと同等 以上であることがわかった。 また、 本非晶質合金の不働態維持電流密度は純 T iと同等に低かった。 これより、 本非晶質合金の不働態皮膜の保護性は、 純 T iと同等以上であることがわかった。 ここで、 本非晶質合金の分極曲線 上には、孔食の発生がみられたことから、合金の耐孔食性を E。 p _{e n}と孔食電 位 （E_{p i t}) の電位差で評価した。 非晶質合金の電位差は SUS 3 1 6 L鋼 での電位差とほぼ同じであったことから、 本非晶質合金の耐孔食性は SU S 31 6 L鋼と同等であると考えられる。

As is clear from the polarization curve shown in FIG. 5, Zr ^ Al ^ CiiA. The amorphous alloy rapidly self-passivated with anodic polarization in the degassed Hanks solution, showed a constant passivation maintaining current, and then pitted. It is considered that the potential of each alloy in a normal body fluid environment is around the natural immersion potential (E. _pen ) in this polarization curve. So E. Polarization resistance corresponding to the corrosion resistance of the material at the _pen Anti ( _Rp ), E. _It was determined from the slope of the current density at the _pen . R _p of the amorphous alloy, an order of magnitude higher than SUS 3 16 L steel, since was several times higher than the pure T i, the corrosion resistance of the amorphous alloys than S US 3 1 6 L Steel It was found to be very high and equal to or higher than pure i. The passivation current density of this amorphous alloy was as low as that of pure Ti. From this, it was found that the passivation film of the present amorphous alloy had a protective property equal to or higher than that of pure Ti. Here, pitting was observed on the polarization curve of the amorphous alloy. It was evaluated by the potential difference between the p _en and pitting power position (E _pit). Since the potential difference of the amorphous alloy was almost the same as that of SUS316L steel, it is considered that the pitting corrosion resistance of this amorphous alloy is equivalent to that of SSUS316L steel.

 Test example 2

Using the _{Z r 5 Q A 1 10 Cu} 40 amorphous alloy, 14 weeks old as an animal model, using the O scan of Uisu evening one (Wistar) based rats were subjected to the following tests.

 (1) Influence of subperiosteal extracorporeal implanted amorphous alloy on living tissue

 In order to examine the clinical application to the plate, histological observations were made by implanting an amorphous alloy ribbon with a length of 10 mm, a width of 2 mm, and a thickness of 0.05 mm into the bone surface and affecting biological tissues. And blood metal concentration (Cu).

 (2) Effect of amorphous alloy inserted into bone marrow on living tissue

 To examine clinical application to intramedullary nails, an amorphous alloy rod with a diameter of 2 mm and a length of 35 mm was inserted into the bone marrow, and for biosafety, blood metal concentration (Cu) and metal content in tissues were included. Evaluation was based on the amount (Cu). For comparison, a group in which Ti-6A1-4V alloy and SUS316L steel were inserted intramedullary and a control group in which no metal was introduced were used as comparison controls.

 (3) Evaluation of in vivo corrosion resistance and durability of amorphous alloys

 The corrosion and wear of the amorphous alloy removed after implantation were evaluated using a scanning electron microscope (SEM) and energy dispersive X-ray spectroscopy (EDS). The metal materials for comparison were Ti-16A1-4V alloy and SUS316L.

(4) Effect of amorphous alloy osteosynthesis material on bone union of fracture Using an amorphous alloy mouth as an intramedullary nail in a femur fracture model, the process of bone formation and bone fusion was quantitatively evaluated. Metal materials for comparison were Ti-6A1-4V alloy and SUS316L.

 The results of Test Example 2 are summarized below.

 (1) Effect of extra-osseous implanted amorphous alloy on living tissue

 The amorphous alloy ribbon was implanted for 6 weeks, but no effect on the living body was observed in the optical microscope image of the surrounding tissue. In addition, among the metal compositions of the amorphous alloy, the concentration of CU in the blood, which could adversely affect the living body, did not increase. From the above, there was no adverse effect on living tissues when the amorphous alloy was implanted into the bone surface for 6 weeks.

 (2) Effect of amorphous alloy inserted into bone marrow on living tissue

 The amorphous alloy alloy was inserted for 6 weeks and 12 weeks, but neither the blood concentration of Cu nor the metal content in the tissue increased compared to the Ti-6A1-4V alloy group and the control group. Was.

 (3) Corrosion resistance and durability of the amorphous alloy in vivo

 The amorphous alloy taken out after embedding was broken on both the ribbon and the rod. No pitting occurred. A qualitative analysis of the metal surface composition with EDS showed that S for SUS316L and P and Ca for Τ Τ-6Α1-4 V alloy were particularly pronounced in the medulla. On the other hand, in the amorphous alloy, Ca appeared slightly in the medulla. The results show that in the medulla, SUS316L corrodes and easily incorporates S as a corrosion product, and Ti-6A1-4V alloy easily forms calcium phosphate on the surface, whereas amorphous This shows that the quality alloy is almost inert.

 (4) Effect of amorphous alloy osteosynthesis material on bone union in fractures

The femoral bone mass was measured by DEXA after a metal material was inserted as an intramedullary nail into a femoral fracture model for 6 weeks (n = 3). The average bone mass ratio compared to the contralateral side (non-operative side) tended to be higher than that of the group using the SUS316L rod and T1-6A1-4V rod for the amorphous alloy. The bone formation at the site of bone fusion after insertion for 12 weeks as an intramedullary nail in a femoral fracture was measured by p QCT (n = 7). Compared to the group using the titanium alloy rod, cortical ossification at the fractured part occurred more actively in the amorphous alloy group, and the bone strength tended to increase.

 From the above results, it was shown that the amorphous alloy is safe as an osteosynthesis material, and is advantageous for bone fusion at a fracture. Industrial applicability

 As described above, the biomaterial member of the present invention does not contain Ni as an allergic source, has excellent corrosion resistance to biological fluids, has high strength, has a low elastic modulus, does not easily react with the living body, and has completed the treatment. It is easy to remove later, and when used as an intramedullary nail etc., it is unlikely to cause so-called stress shield ゃ pulmonary infarction.

 Therefore, dental crowns, inlays, crowns, prostheses, and human roots, such as orthopedic spinal fixation materials such as intramedullary nails and osteosynthesis plates, fracture fixation materials, human joints, and intervertebral spacers It can be suitably used as a biomaterial member useful for an orthodontic wire, a surgical instrument for general surgery, and the like.

Claims

The scope of the claims 1. Degassed An amorphous alloy whose polarization resistance in Hanks liquid is 4 x 10 ⁶ Ωcm ² or more and whose immersion potential-pitting corrosion potential window is 0.25 V or more Biomaterials.  2. The biomaterial member according to claim 1, comprising an amorphous alloy having a Young's modulus of 100 GPa or less.  3. The biomaterial member according to claim 1, wherein the amorphous alloy has a composition represented by one of the following general formulas (1) and (2).  General formula (1): Μ ^ Μ ^ Μ '^ Μ ^ Μ ^ (However, M ¹ is one, two or three elements selected from Zr, Hf and Ti; M ² is Cu, Fe, Co, Mn, Nb, V, Cr, Zn, at least one element a l, at least one element selected from the group consisting of Sn and G a, M ³ is B, C ;, N, P, selected from the group consisting of S i and 0, M ⁴ is T a, at least one element selected from the group consisting of W and Mo, M ⁵ Ho Au, P t, at least one element selected from the group consisting of P d and a g, a, b, c , d , And e are in atomic% respectively, 25 ≤ a ≤ 85, 15 ≤ b ≤ 75 0 ≤ c ≤ 30, 0 ≤ d ≤ 15, 5, 0 ≤ e ≤ 15.  General formula (2): A 1: l to 10%, Ga: 0.5 to 4%, P: 9 to 15%, C: 5 to 7%, B: 2 to: L 0% in atomic% Si: 0 to: L 5%, Ge: 0 to 4%, Fe: Fe-based amorphous alloy having the remaining composition.  4. The biomaterial according to any one of claims 1 to 3, wherein the biomaterial is at least one of an intramedullary nail, an osteosynthesis plate, an interlocking nail, and a bone fixing screw.