JP2005029845A - Titanium alloy - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a new titanium alloy having a low melting point while keeping high performances equal to those of conventional titanium alloys with respect to corrosion resistance, biocompatibility, etc. <P>SOLUTION: The titanium alloy contains Nb, Zr, at least one element selected from the group consisting of Cr, Fe and Si, and the balance Ti with inevitable impurities. Further, the titanium alloy contains, by mass, 25 to 35% Nb, 5 to 20% Zr, ≥0.5% of at least one element selected from the group consisting of Cr, Fe and Si, and the balance Ti with inevitable impurities. <P>COPYRIGHT: (C)2005,JPO&NCIPI

Description

【００１２】
〔チタン合金の作製〕
各設計合金は、総重量４５ｇとなるように秤量後、非消耗電極式アーク炉を用いて高純度アルゴンガス雰囲気中にて溶製を行った。このとき、溶製によるミクロ偏析等を避けるため、溶製させたボタンインゴットを上下逆にし再溶製するフィリッピング処理を５回以上行った。
【００１３】
〔鋳型の作製〕
鋳型の作製に用いる埋没材には、高融点高活性材料であるチタン合金との反応を軽減する目的で、市販のマグネシア系埋没材からシリカ系の膨張材およびアルミナを除去し、主基材であるマグネシアを微細化させた改良型埋没材（商品名：タイタンベストＣ改２：岡崎鉱産物 （株） 製）を用いた。蒸留水および専用液を４：１の割合で混合した混合液を作製し、埋没材およびその混合液を１００：１８の割合で混合し、真空撹拌機を用いて１分間真空撹拌を行った。撹拌後のスラリーを型枠へバイブレーターで振動を与えながら流し込み、室温（２９５Ｋ）の大気中で２時間保持し乾燥させた。マグネシア系埋没材の焼成は、室温から０．０８３Ｋ／ｓで加熱し、１３７３Ｋで３．６ｋｓ保持後、室温まで炉冷し、鋳造に供した。
【００１４】
〔鋳造工程〕
チタン合金の鋳造には、真空加圧式鋳造機（商品名：オートキャストＨＣ‐ＩＩＩ（Ｇ．Ｃ．（株）製））を使用した。上記［表１］に示したＡ〜Ｅの設計合金を、真空（１０．２×１０^−６Ｐａ）に保持したチャンバー内にて、３００Ａの直流アークにより加熱した。加熱により合金の形状がドーム状に変形後８秒間保持し、アルゴンガス加圧（０．８９ＭＰａ）にて鋳造を行った。鋳造体は、大気中にて鋳型とともに室温まで冷却した。
【００１５】
〔融点測定〕
Ａ〜Ｅの各設計合金の融点測定には、約２５００Ｋまで温度測定可能な直径０．４ｍｍのＷ５Ｒｅ−Ｗ２６Ｒｅ熱電対を用いた。鋳型内にこの熱電対を設置してから各設計合金の鋳造を行い、溶湯状態からの温度変化を測定し、各設計合金の冷却曲線を作製した。各設計合金の融点は、冷却時の相変態に起因する冷却曲線の遷移温度あるいは過冷後の飽和温度とした。
【００１６】
〔試験片作製〕
各設計合金の鋳造体における引張特性を評価するため、図１に示す標点間部直径φ３．５ｍｍ引張試験片を歯科精密鋳造にて作製した。すなわち、図１に示す引張試験片のワックスパターンをポリエチレンおよびワックスにて作製し、湯道および湯口を装着後、前述した埋没材を用いて鋳型を作製し、各設計合金の鋳造に供した。各設計合金を鋳造後、鋳型から試験片を取り出しサンドブラストにて試験片表面を研掃し、機械加工にて湯口部を切断して引張試験に供した。
【００１７】
〔引張試験〕
引張試験は、容量９．８ｋＮのインストロン型試験機を用い、クロスヘッド速度８．３３×１０^−６ｍ／ｓで室温の大気中にて行った。荷重の測定は、試験機のロードセルにより行い、変位の測定は、試験片の標点間部に直接取り付けた歪ゲージおよび読取顕微鏡にて行った。
【００１８】
〔ビッカース硬さ試験〕
ビッカース硬さ試験は、引張試験と同形状の試験片のつかみ部断面を♯１５００までのエメリー紙により湿式研磨し、湿式研磨後の試験片表面（鋳肌）近傍からその内部にかけて２０μｍの間隔で深さ５００μｍまで測定を行った。測定は、マイクロビッカース硬さ試験機を用いて、荷重２００ｇおよび保持時間１５ｓにて行った。
【００１９】
〔ミクロ組織観察〕
ミクロ組織観察は、ビッカース硬さ試験後の試験片を二酸化シリコン懸濁液によるバフ研磨によって鏡面に仕上げ、その後５％フッ酸水溶液により試験片表面の腐食を行い、光学顕微鏡を用いて行った。
【００２０】
〔破面観察〕
引張試験後の破面観察には、走査型電子顕微鏡（ＳＥＭ）を用いた。また、破面観察用試料は、引張試験片の破断面から約３ｍｍの部位を機械加工により切断し、アセトンによる超音波洗浄を行った。
【００２１】
これらの試験によって、以下の結果が得られた。
〔設計合金の融点〕
Ａ〜Ｅの各設計合金およびＴＮＴＺ合金の鋳型内における溶湯の凝固過程の冷却曲線を図２に示す。
図２に示すように、各設計合金の冷却初期に現れる冷却速度遷移点は、従来のチタン合金であるＴＮＴＺ合金のそれと比較して全般的に低下した。また、各設計合金の融点の値は、ＴＮＴＺ合金の融点の値よりも約５０Ｋ〜３７０Ｋの範囲で低い値を示した。その中でも、Ａ：Ｔｉ−２９Ｎｂ−１３Ｚｒ−２Ｃｒ合金においては、約２０５０Ｋと最も低い融点を示した。
これらの結果より、高融点の金属元素であるＴａを排除し、低融点化元素であるＣｒ、Ｓｉ、Ｆｅのうち少なくとも１種の元素を添加することによって、例えば従来のチタン合金であるＴＮＴＺ合金よりも低融点を有するチタン合金を実現できることが見出された。
【００２２】
〔鋳造体の外観〕
Ａ〜Ｅの各設計合金、及び、比較例としてＴＮＴＺ合金の鋳造体の外観を目視にて比較した。その結果、ＴＮＴＺ合金では、鋳造体表面が黒鉛色を呈し、鋳型材の付着傾向が高いことが判明した。これに対して、Ａ〜Ｅの各設計合金の鋳造体表面は、ゴールド色を呈する領域を一部に有しており、全体的に金属光沢を呈した。
チタン合金の表面には、一般に安定な不動態化層である酸化皮膜が形成されており、その酸化被膜の厚さによって、合金表面の色彩が変化することが知られている。Ａ〜Ｅの設計合金表面で観察されたゴールド色は、約２０ｎｍの酸化皮膜厚さに起因していると考えられる。また、他の領域では金属光沢を有していることから、各設計合金の融点がＴＮＴＺ合金の融点よりも低下し、鋳型との反応性が低下したものと考えられる。
【００２３】
〔鋳造体断面のビッカース硬さ〕
Ａ〜Ｅの各設計合金、及び、比較例としてＴＮＴＺ合金の引張試験片つかみ部断面における試験片表面（鋳肌）近傍から内部（表面から約５００μｍ）にかけてのビッカース硬さの分布を図３に示す。
図３に示すように、各設計合金の鋳肌表面から内部にかけての約２００μｍの領域には、試験片中心部付近のマトリックスよりもビッカース硬さの値が高くなっている領域が存在する。この領域は、鋳造時における鋳型との反応硬化層と考えられ、その厚さは、ＴＮＴＺ合金における反応硬化層の厚さとほぼ同等である。Ａ〜Ｅの各設計合金の試験片表面近傍のビッカース硬さは４００Ｈｖから５００Ｈｖの範囲内に分布しているのに対して、ＴＮＴＺ合金の試験片表面のビッカース硬さは約５６０Ｈｖと最も高い値を示している。これは、各設計合金の融点がＴＮＴＺ合金のそれより低いため、鋳型との反応性が低下し、酸素の拡散が抑制されたことに起因すると考えられる。
【００２４】
〔引張特性〕
図４は、Ａ〜Ｅの各設計合金及びＴＮＴＺ合金の引張り強さ及び伸び量の試験結果を示している。
図４に示すように、Ｄ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｃｒ−０．５Ｆｅ合金を除く全ての設計合金の引張強さは７００ＭＰａを超えており、ＴＮＴＺ合金の引張強さよりも約１００ＭＰａ上昇する傾向を示した。
また、図４に示すように、Ａ：Ｔｉ−２９Ｎｂ−１３Ｚｒ−２Ｃｒ合金、Ｂ：Ｔｉ−２９Ｎｂ−１５Ｚｒ−１．５Ｆｅ合金およびＤ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｃｒ−０．５Ｆｅ合金の伸び量は、一般に歯科製品に必要であると言われている１０％以上の伸び量を示した。その中でも、Ｄ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｃｒ−０．５Ｆｅ合金の伸び量は２１％と最も高い値を示しており、ＴＮＴＺ合金の伸び（約２３％）とほぼ同等の値であり高い延性を示した。
その一方、Ｓｉを含有するＣ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｓｉ合金およびＥ：Ｔｉ−２９Ｎｂ−１８Ｚｒ−２Ｃｒ−０．５Ｓｉ合金の伸び量は３％以下の値であり、その他の合金と比較すると脆化する傾向を示した。
これらの結果より、本発明のチタン合金、特に、Ａ：Ｔｉ−２９Ｎｂ−１３Ｚｒ−２Ｃｒ合金およびＤ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｃｒ−０．５Ｆｅ合金の強度延性バランスは、従来の生体用チタン合金であるＴＮＴＺ合金よりも良好であることが判明した。
【００２５】
〔引張試験後の破断部形態〕
Ａ〜Ｅの各設計合金、及び、比較例としてＴＮＴＺ合金の引張試験後における破断部の電子顕微鏡写真を図５に示す。引張試験において伸びが１０％以上であったＡ：Ｔｉ−２９Ｎｂ−１３Ｚｒ−２Ｃｒ合金、Ｂ：Ｔｉ−２９Ｎｂ−１５Ｚｒ−１．５Ｆｅ合金およびＤ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｃｒ−０．５Ｆｅ合金では５１％以上の断面減少率を有するのに対して、Ｃ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｓｉ合金およびＥ：Ｔｉ−２９Ｎｂ−１８Ｚｒ−２Ｃｒ−０．５Ｓｉ合金では、破面近傍にくびれは確認されず、応力軸に対して垂直にき裂が進展する低延性的な破壊挙動を示した。
【００２６】
〔鋳造体のミクロ組織〕
Ａ〜Ｅの各設計合金およびＴＮＴＺ合金の試験片表面近傍および内部におけるミクロ組織の光学顕微鏡写真を図６に示す。引張試験における伸びが最も良好であったＤ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｃｒ−０．５Ｆｅ合金の試験片表面近傍の結晶粒径は、他の設計合金のそれと比較して若干粗大化傾向を示した。
また、図６に示すように、低延性を示したＣ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｓｉ合金およびＥ：Ｔｉ−２９Ｎｂ−１８Ｚｒ−２Ｃｒ−０．５Ｓｉ合金の試験片表面の結晶粒径は、内部の結晶粒径と比較して小さくなる傾向を示している。これに対し、Ａ、Ｂ、Ｄの設計合金の試験片表面の結晶粒径は、試験片内部における結晶粒径とほぼ同等になる傾向を示している。
よって、Ｃ：Ｔｉ−２９Ｎｂ−１０Ｚｒ−０．５Ｓｉ合金およびＥ：Ｔｉ−２９Ｎｂ−１８Ｚｒ−２Ｃｒ−０．５Ｓｉ合金のように試験片表面と内部との結晶粒径に相違が生じる場合、鋳造体の延性を低下させると考えられる。また、両合金では合金添加元素であるＳｉの化合物が形成され、そのＳｉの化合物が延性を低下させていることが示唆される。
【００２７】
以上の結果より、本願発明のチタン合金が極めて有用であることを確認することができた。
すなわち、Ｎｂ、Ｚｒを含有し、さらに、Ｃｒ、Ｆｅ、Ｓｉからなる群より選択される少なくとも１種の元素を含有するとともに、残部がＴｉ及び不可避的不純物からなる新しいチタン合金を開発した。このようなチタン合金を鋳造することで試験片を製造し、従来のチタン合金（例えばＴＮＴＺ合金）との比較のための各種試験を行った。これにより、本願発明のチタン合金は、従来のチタン合金とほぼ同等の引張り強度及び延性を有することを確認するとともに、従来のチタン合金よりも低い融点を有することを確認することができた。
本願発明のチタン合金は、生体適合性が高いために歯科材料あるいは生体組織代替材料としての有用性が極めて大きい。また、本願発明のチタン合金は、融点が低いために鋳型との反応性が小さく、鋳造により製造される義歯等の精密歯科材料として極めて好適に使用することができる。
本願発明のチタン合金は、前述した用途以外にも、インプラント材、人工関節、歯科矯正材等の種々の生体組織用代替材またはその一部の補助材として使用することも可能である。
【００２８】
また、本願発明のチタン合金の組成は、上記［表１］に示したＡ〜Ｅのチタン合金の組成に限るものではない。チタン合金の融点を低下させるためには、Ｃｒ、Ｆｅ、Ｓｉからなる群より選択される少なくとも一種の元素を添加すればよく、例えば、Ｆｅ及びＳｉを添加してもよいし、Ｃｒ、Ｆｅ、Ｓｉのすべてを添加してもよい。
【００２９】
【発明の効果】
以上説明したように、本願発明によれば、耐食性、人体適合性などの点において従来のチタン合金並みの高い性能を維持しながらも、低融点を有する新しいチタン合金を提供することができた。
【図面の簡単な説明】
【図１】Ａ〜Ｅの各設計合金およびＴＮＴＺ合金の引張特性を評価するための引張試験片の形状を示す図である。
【図２】Ａ〜Ｅの各設計合金およびＴＮＴＺ合金の鋳型内における溶湯の凝固過程の冷却曲線を示す図である。
【図３】Ａ〜Ｅの各設計合金およびＴＮＴＺ合金の引張試験片つかみ部断面における試験片表面（鋳肌）近傍から内部（表面から約５００μｍ）にかけてのビッカース硬さの分布を示す図である。
【図４】Ａ〜Ｅの各設計合金及びＴＮＴＺ合金の引張り強さ及び伸び量の試験結果を示す図である。
【図５】Ａ〜Ｅの各設計合金及びＴＮＴＺ合金の引張試験後における破断部の電子顕微鏡写真を示している。
【図６】Ａ〜Ｅの各設計合金およびＴＮＴＺ合金の試験片表面近傍および内部におけるミクロ組織の光学顕微鏡写真を示している。[0001]
BACKGROUND OF THE INVENTION
The present invention relates to a titanium alloy, and more particularly to a biomedical titanium alloy that can be suitably used as a dental material or a medical material.
[0002]
[Prior art]
Conventionally, as a biomedical titanium alloy used for a dental artificial tooth root or a medical artificial bone material, for example, a biomedical titanium alloy disclosed in [Patent Document 1] is known.
According to the biomedical titanium alloy disclosed in [Patent Document 1], Nb and Ta are contained in a total of 20 wt% to 60 wt%, Zr is contained as necessary, and the balance is composed of Ti and inevitable impurities. ing. According to this titanium alloy, it has high corrosion resistance, does not contain V (vanadium), which has been pointed out to be toxic to the human body, contains Nb and Ta having high biocompatibility, and is a β-type titanium alloy. A moderate elastic modulus close to that of bone can be obtained. Such a titanium alloy for living organisms has been specifically proposed as a Ti-29Nb-13Ta-4.6Zr alloy (hereinafter referred to as a TNTZ alloy) by the inventors of the present application.
[0003]
[Patent Document 1]
Japanese Patent Laid-Open No. 10-219375
[Problems to be solved by the invention]
However, since the conventional titanium alloy (TNTZ alloy) described above contains a high melting point element such as Nb and Ta, the melting point is extremely higher than that of a conventional biomaterial. Therefore, when considering the application in the dental field, etc., where casting products are the majority, defects such as seizure on the surface of this precision cast alloy and formation of a surface reaction layer due to reaction with the mold occur. It was feared that it would greatly affect the mechanical properties.
[0005]
The present invention has been made in view of such problems, and its object is to maintain a low melting point while maintaining the same high performance as conventional titanium alloys in terms of corrosion resistance and human compatibility. It is to provide a new titanium alloy.
[0006]
[Means for Solving the Problems]
In order to solve the above-described problems, the inventors of the present application exclude Ta having a high melting point from the composition of a conventional titanium alloy (for example, a TNTZ alloy), and newly lower the melting point of Cr, Si, Fe, etc. It has been found that elements may be added.
That is, the first invention of the present application contains Nb and Zr, further contains at least one element selected from the group consisting of Cr, Fe, and Si, and the balance is made of Ti and inevitable impurities. Titanium alloy. According to such a titanium alloy, it does not contain an element that is pointed out to be toxic or allergenic, such as V (vanadium), and has high biocompatibility and corrosion resistance by containing Nb and Zr. A titanium alloy having a lower melting point than an alloy (for example, a TNTZ alloy) can be realized.
[0007]
Moreover, 2nd invention of this application contains Nb: 25-35% and Zr: 5-20% by mass ratio in the said 1st invention, Furthermore, it selects from the group which consists of Cr, Fe, and Si The titanium alloy contains 0.5% or more of at least one element that is formed and the balance is Ti and inevitable impurities. The reason why the content of each element is set in such a range is because of the following advantages.
That is, the Nb content is preferably 25% to 35% in terms of mass ratio with respect to the entire titanium alloy. This is because when the Nb content is smaller than 25%, the tendency of the α phase to precipitate in the alloy structure becomes strong, and when the Nb content is larger than 35%, the tendency of the alloy to become insufficient becomes strong. It is.
The content of Zr is preferably 5 to 20% by mass ratio with respect to the entire titanium alloy. This is because when the Zr content is set within this range, a titanium alloy having high biocompatibility can be realized in addition to an increase in stability of the titanium alloy.
The content of at least one element selected from the group consisting of Cr, Fe, and Si is preferably 0.5% or more in terms of a mass ratio with respect to the entire titanium alloy. This is because if these elements are contained at least 0.5%, the melting point of the titanium alloy can be lowered.
[0008]
A third invention of the present application is a titanium alloy for living organisms used as a dental material or a biological tissue substitute material in the first invention or the second invention.
Since the titanium alloy of the present invention has a low melting point, the reactivity with the mold is reduced during casting. Therefore, the usefulness as a titanium alloy for living bodies is great, and it can be suitably used as a precision dental material such as artificial tooth roots and artificial teeth manufactured by casting, or a substitute material for living tissues such as artificial bones, artificial limbs and artificial legs.
[0009]
DETAILED DESCRIPTION OF THE INVENTION
Hereinafter, embodiments of the present invention will be described in detail.
In the present embodiment, a titanium alloy having a low melting point has been developed while maintaining various performances such as corrosion resistance and biocompatibility of conventional titanium alloys. In order to evaluate such a titanium alloy, five types of titanium alloys were cast, and the tensile properties and hardness of the cast titanium alloys were evaluated.
[0010]
[Determination of titanium alloy constituent elements]
The constituent elements of Ti-29Nb-13Ta-4.6Zr alloy (TNTZ alloy), which is a conventional titanium alloy for living organisms, are determined in consideration of factors such as cytotoxicity, biological tissue compatibility and polarization resistance of various pure metals. Has been. Therefore, in the present invention, the Ti—Nb—Zr alloy in which Ta having a high melting point is excluded from the TNTZ alloy is adopted as the basic composition in consideration of biocompatibility. Further, according to the binary phase diagram, at least one element selected from Cr, Si and Fe was selected as an alloy additive element as an element for effectively lowering the melting point. The alloy design uses a d-electron alloy design method in which the strength of the bond with the alloy element is evaluated using two alloy parameters of bond order (Bo value) and d-orbital energy level (Md value). It was. Thus, as an example of the titanium alloy of the present invention, five types of β-type titanium alloys A to E shown in the following [Table 1] were designed, and these titanium alloys were produced and evaluated.
[0011]
[Table 1]

[0012]
[Production of titanium alloy]
Each design alloy was weighed so as to have a total weight of 45 g, and then melted in a high-purity argon gas atmosphere using a non-consumable electrode type arc furnace. At this time, in order to avoid microsegregation or the like due to melting, the filing process of remelting the melted button ingot upside down was performed five times or more.
[0013]
[Production of mold]
In order to reduce the reaction with the titanium alloy, which is a high melting point and high activity material, the silica base expansion material and alumina were removed from the commercially available magnesia base investment material, and the main base material An improved investment material (trade name: Titan Best C 2: manufactured by Okazaki Mineral Co., Ltd.) in which a certain magnesia was refined was used. A mixed liquid in which distilled water and a dedicated liquid were mixed at a ratio of 4: 1 was prepared, the investment material and the mixed liquid were mixed at a ratio of 100: 18, and the mixture was vacuum stirred using a vacuum stirrer for 1 minute. The agitated slurry was poured into the mold while being vibrated with a vibrator, and kept in the atmosphere at room temperature (295 K) for 2 hours to be dried. The magnesia investment was fired from room temperature at 0.083 K / s, held at 1373 K for 3.6 ks, cooled to room temperature, and subjected to casting.
[0014]
[Casting process]
A vacuum pressure casting machine (trade name: Autocast HC-III (manufactured by GC Corporation)) was used for casting of the titanium alloy. The design alloys A to E shown in the above [Table 1] were heated by a 300 A DC arc in a chamber maintained in a vacuum (10.2 × 10 ⁻⁶ Pa). The shape of the alloy was changed to a dome shape by heating and held for 8 seconds, and casting was performed with argon gas pressurization (0.89 MPa). The casting was cooled to room temperature together with the mold in the atmosphere.
[0015]
[Melting point measurement]
A W5Re-W26Re thermocouple having a diameter of 0.4 mm capable of measuring the temperature up to about 2500 K was used to measure the melting point of each of the designed alloys A to E. After the thermocouple was installed in the mold, each designed alloy was cast, the temperature change from the molten state was measured, and the cooling curve of each designed alloy was prepared. The melting point of each design alloy was the transition temperature of the cooling curve caused by the phase transformation during cooling or the saturation temperature after supercooling.
[0016]
[Test specimen preparation]
In order to evaluate the tensile characteristics in the cast body of each design alloy, a tensile test piece having a diameter of 3.5 mm between the gauge points shown in FIG. 1 was prepared by dental precision casting. That is, the wax pattern of the tensile test piece shown in FIG. 1 was made of polyethylene and wax, and after the runner and the sprue were attached, a mold was made using the above-mentioned investment material and used for casting of each design alloy. After casting each design alloy, the test piece was taken out from the mold, the surface of the test piece was polished by sandblasting, and the gate was cut by machining to be used for a tensile test.
[0017]
[Tensile test]
The tensile test was performed in an atmosphere at room temperature using an Instron type testing machine having a capacity of 9.8 kN and a crosshead speed of 8.33 × 10 ⁻⁶ m / s. The load was measured with a load cell of a testing machine, and the displacement was measured with a strain gauge and a reading microscope attached directly between the gauge points of the test piece.
[0018]
[Vickers hardness test]
In the Vickers hardness test, the cross section of the grip part of the test piece of the same shape as the tensile test is wet-polished with emery paper up to # 1500, and the distance from the surface of the test piece (cast surface) after wet polishing to the inside is 20 μm. Measurements were made to a depth of 500 μm. The measurement was performed using a micro Vickers hardness tester with a load of 200 g and a holding time of 15 s.
[0019]
[Microstructure observation]
Microscopic observation was performed using an optical microscope by finishing the specimen after the Vickers hardness test to a mirror surface by buffing with a silicon dioxide suspension and then corroding the specimen surface with a 5% hydrofluoric acid aqueous solution.
[0020]
[Fracture surface observation]
A scanning electron microscope (SEM) was used for fracture surface observation after the tensile test. Further, the fracture surface observation sample was obtained by cutting a portion of about 3 mm from the fracture surface of the tensile test piece by machining and performing ultrasonic cleaning with acetone.
[0021]
These tests yielded the following results:
[Melting point of design alloy]
FIG. 2 shows cooling curves of the solidification process of the molten metal in the molds of each of the designed alloys A to E and the TNTZ alloy.
As shown in FIG. 2, the cooling rate transition point appearing at the initial stage of cooling of each designed alloy was generally lower than that of the TNTZ alloy, which is a conventional titanium alloy. Moreover, the value of the melting point of each design alloy showed a lower value in the range of about 50K to 370K than the melting point value of the TNTZ alloy. Among them, the A: Ti-29Nb-13Zr-2Cr alloy showed the lowest melting point of about 2050K.
From these results, Ta, which is a high melting point metal element, is eliminated, and at least one element of Cr, Si, Fe, which is a low melting point element, is added, for example, a conventional titanium alloy, a TNTZ alloy It has been found that a titanium alloy having a lower melting point can be realized.
[0022]
[Appearance of cast body]
The appearance of each of the designed alloys A to E and a cast of a TNTZ alloy as a comparative example were visually compared. As a result, it was found that in the TNTZ alloy, the casting surface had a graphite color, and the mold material had a high tendency to adhere. On the other hand, the cast body surface of each of the designed alloys A to E had a region exhibiting a gold color in part and exhibited a metallic luster as a whole.
It is known that an oxide film, which is a stable passivation layer, is generally formed on the surface of the titanium alloy, and the color of the alloy surface changes depending on the thickness of the oxide film. The gold color observed on the surface of the designed alloys A to E is considered to be due to the oxide film thickness of about 20 nm. Moreover, since it has a metallic luster in other regions, it is considered that the melting point of each design alloy is lower than the melting point of the TNTZ alloy and the reactivity with the mold is lowered.
[0023]
[Vickers hardness of the casting cross section]
FIG. 3 shows the distribution of Vickers hardness from the vicinity of the test piece surface (cast surface) to the inside (about 500 μm from the surface) in the cross section of the tensile test piece gripping part of each of the designed alloys A to E and the TNTZ alloy as a comparative example. Show.
As shown in FIG. 3, there is a region where the value of Vickers hardness is higher than the matrix near the center of the test piece in the region of about 200 μm from the casting surface to the inside of each design alloy. This region is considered to be a reaction hardened layer with the mold at the time of casting, and its thickness is almost equal to the thickness of the reaction hardened layer in the TNTZ alloy. The Vickers hardness in the vicinity of the specimen surface of each design alloy of A to E is distributed in the range of 400 Hv to 500 Hv, whereas the Vickers hardness of the specimen specimen of TNTZ alloy is the highest value of about 560 Hv. Is shown. This is considered to be due to the fact that the melting point of each design alloy is lower than that of the TNTZ alloy, so that the reactivity with the mold is lowered and the diffusion of oxygen is suppressed.
[0024]
(Tensile properties)
FIG. 4 shows the test results of the tensile strength and elongation amount of each of the designed alloys A to E and the TNTZ alloy.
As shown in FIG. 4, the tensile strength of all the design alloys except D: Ti-29Nb-10Zr-0.5Cr-0.5Fe alloy exceeds 700 MPa, which is about 100 MPa higher than the tensile strength of the TNTZ alloy. Showed a tendency to
Further, as shown in FIG. 4, A: Ti-29Nb-13Zr-2Cr alloy, B: Ti-29Nb-15Zr-1.5Fe alloy and D: Ti-29Nb-10Zr-0.5Cr-0.5Fe alloy. The elongation amount showed an elongation amount of 10% or more, which is generally said to be necessary for dental products. Among them, the elongation amount of the D: Ti-29Nb-10Zr-0.5Cr-0.5Fe alloy shows the highest value of 21%, which is almost equal to the elongation of the TNTZ alloy (about 23%). High ductility was exhibited.
On the other hand, the elongation of the C: Ti-29Nb-10Zr-0.5Si alloy containing Si and the E: Ti-29Nb-18Zr-2Cr-0.5Si alloy is 3% or less. In comparison, it showed a tendency to become brittle.
From these results, the strength ductility balance of the titanium alloys of the present invention, in particular, A: Ti-29Nb-13Zr-2Cr alloy and D: Ti-29Nb-10Zr-0.5Cr-0.5Fe alloy is It was found to be better than a TNTZ alloy which is a titanium alloy.
[0025]
[Fracture form after tensile test]
The electron micrograph of the fracture | rupture part after the tensile test of each design alloy of AE and a TNTZ alloy as a comparative example is shown in FIG. A: Ti-29Nb-13Zr-2Cr alloy, B: Ti-29Nb-15Zr-1.5Fe alloy, and D: Ti-29Nb-10Zr-0.5Cr-0.5Fe, whose elongation was 10% or more in the tensile test The alloy has a cross-sectional reduction ratio of 51% or more, whereas the C: Ti-29Nb-10Zr-0.5Si alloy and the E: Ti-29Nb-18Zr-2Cr-0.5Si alloy are constricted near the fracture surface. No low ductility fracture behavior was observed in which cracks propagated perpendicular to the stress axis.
[0026]
[Microstructure of cast body]
FIG. 6 shows optical micrographs of the microstructures near and inside the specimen surface of each of the designed alloys A to E and the TNTZ alloy. The crystal grain size in the vicinity of the test piece surface of D: Ti-29Nb-10Zr-0.5Cr-0.5Fe alloy, which had the best elongation in the tensile test, was slightly coarser than that of other designed alloys. showed that.
Further, as shown in FIG. 6, the crystal grain sizes of the test piece surfaces of C: Ti-29Nb-10Zr-0.5Si alloy and E: Ti-29Nb-18Zr-2Cr-0.5Si alloy exhibiting low ductility are as follows. It shows a tendency to be smaller than the internal crystal grain size. On the other hand, the crystal grain size on the test piece surface of the design alloys of A, B, and D tends to be almost equal to the crystal grain size inside the test piece.
Therefore, when there is a difference in the crystal grain size between the test piece surface and the inside, such as C: Ti-29Nb-10Zr-0.5Si alloy and E: Ti-29Nb-18Zr-2Cr-0.5Si alloy, This is considered to reduce the ductility of the steel. Further, in both alloys, a compound of Si as an alloying element is formed, which suggests that the compound of Si decreases the ductility.
[0027]
From the above results, it was confirmed that the titanium alloy of the present invention was extremely useful.
That is, a new titanium alloy containing Nb and Zr, further containing at least one element selected from the group consisting of Cr, Fe, and Si and the balance being Ti and inevitable impurities has been developed. A test piece was manufactured by casting such a titanium alloy, and various tests for comparison with a conventional titanium alloy (for example, a TNTZ alloy) were performed. Thereby, while confirming that the titanium alloy of this invention has the tensile strength and ductility substantially equivalent to the conventional titanium alloy, it has confirmed that it had melting | fusing point lower than the conventional titanium alloy.
Since the titanium alloy of the present invention has high biocompatibility, it is extremely useful as a dental material or a biological tissue substitute material. In addition, the titanium alloy of the present invention has a low melting point and thus has low reactivity with the mold, and can be used very suitably as a precision dental material such as a denture manufactured by casting.
The titanium alloy of the present invention can be used as various substitute materials for living tissue such as implant materials, artificial joints, orthodontic materials, or a part of auxiliary materials in addition to the above-described uses.
[0028]
The composition of the titanium alloy of the present invention is not limited to the composition of the titanium alloys A to E shown in the above [Table 1]. In order to lower the melting point of the titanium alloy, at least one element selected from the group consisting of Cr, Fe, and Si may be added. For example, Fe and Si may be added, or Cr, Fe, All of Si may be added.
[0029]
【The invention's effect】
As described above, according to the present invention, it was possible to provide a new titanium alloy having a low melting point while maintaining the same high performance as conventional titanium alloys in terms of corrosion resistance and human body compatibility.
[Brief description of the drawings]
FIG. 1 is a diagram showing the shape of a tensile test piece for evaluating the tensile properties of each of the designed alloys A to E and a TNTZ alloy.
FIG. 2 is a diagram showing a cooling curve of a solidification process of a molten metal in a mold of each of the designed alloys A to E and a TNTZ alloy.
FIG. 3 is a diagram showing the distribution of Vickers hardness from the vicinity of the test piece surface (cast surface) to the inside (about 500 μm from the surface) in the cross section of the tensile test piece gripping part of each of the designed alloys of A to E and the TNTZ alloy. .
FIG. 4 is a diagram showing the test results of tensile strength and elongation of each design alloy of A to E and a TNTZ alloy.
FIG. 5 shows electron micrographs of fracture parts after tensile tests of each of the designed alloys A to E and the TNTZ alloy.
FIG. 6 shows optical micrographs of microstructures in and near the test piece surface of each of the designed alloys AE and the TNTZ alloy.

Claims (3)

A titanium alloy containing Nb, Zr, and further containing at least one element selected from the group consisting of Cr, Fe, and Si, with the balance being Ti and inevitable impurities. The titanium alloy according to claim 1,
In terms of mass ratio, Nb: 25 to 35%, Zr: 5 to 20% are contained, and at least one element selected from the group consisting of Cr, Fe, and Si is further contained in total of 0.5% or more. In addition, a titanium alloy with the balance being Ti and inevitable impurities. The titanium alloy according to claim 1 or 2,
Biomedical titanium alloy used as a dental material or a substitute for living tissue.

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