CA2110779A1

CA2110779A1 - Medical implants of biocompatible low modulus titanium alloy

Info

Publication number: CA2110779A1
Application number: CA002110779A
Authority: CA
Inventors: James A. Davidson; Ajit K. Mishra; Paul Kovacs
Original assignee: Individual
Current assignee: Smith and Nephew Inc
Priority date: 1992-12-07
Filing date: 1993-12-06
Publication date: 1994-06-08
Also published as: US5716400A; AU5219693A; US5690670A; JPH06233811A; US5782910A; EP0601804A1; US5562730A; US5477864A; US5713947A; US5676632A; US5685306A

Abstract

ABSTRACT

Biocompatible medical implants from a high strength titanium alloy with low elastic modulus containing titanium, about 10-20 wt% or 35 to about 50 wt% niobium and up to 20 wt% zirconium. In particular, the titanium alloy has a modulus of elasticity closer to that of bone than other typically used metal alloys and does not include any elements which have been shown or suggested as having short or long term potential adverse effects from a standpoint of biocompatibility.

Description

~.

MEDICA~ I~PLAN~8 OF BIOCOMPA~IBLE LO~
MODULUB ~I~ANIUN ~LLOY
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This inv~ntion relates to high strength, biocompa~ible metallic implants. In particular, the invention is of titanium alloy medical implants that have a low modulus of elasticity and high strength 5. produced by a series of specific metallurgical steps to ~ -which the alloy is subjected. Further, the alloy does nok include any elements which have been shown or suggested as having short term or long term potential adverse effects when implanted in the human body. ~
10. :
For many applications there has been, and there continues to be, a need for a metal that has a low modulus of elasticity but is alsio strong, fatigue-resistant, corrosion resistant, and has a hard surface 15. that is resistant to abrasive wear. For instance, in the orthopaedic implant art, metals are still the most commonly used material for fahricating load-bearing -implants such as, for instanc~!, hip joints and knee joints. -20.
Metals and metal alloys such as stainless steel, vitalium (cobalt alloy) and ti.tanium have been used successfully. These materials have the requisite ~:
strength characteristics but typically have not been 25. resilient or flexible enough to form an optimum implant material. Also, many alloys contain elements such as aluminium, vanadium, cobalt, nickel, molybdenum, and chromium which recent studies have suggested might have some long term adverse effects on human patients.
30~
Many of the metal alloys typically used in prosthetic implants were developed for other applications, such as Ti-6A1-4V alloy in the aircraft -;: 2 ~1~77~

industry. These alloys were later thought to be suitable ~or use as implant materials because they possess mechanical strength and appeared to have acceptable levels of biocompatibility. However, these ;. metals typically have elastic moduli much higher than that of bone, for example, 316 stainless steel has an elastic modulus of about 200 GPa while that of cast heat-treated Co-Cr-Mo alloy is about 240 GPa.

10. It has also been found that many of these metals will corrode to some extent in body fluids thereby releasing ions ~hat might possibly be harmful over a prolonged period of time. It is now believed that the corrosive effects of body fluids is due both to 15. chemical and electro-chemical processes, with corrosion products forming when certain commonly-used metal ;~
alloys ionize from corrosion processes in the body.
For example, aluminium metal ions have been associated with Alzheimer's disease and vanadium, cobalt, 20. molybdenum, nickel and chromium are suspected of being toxic or carcinogenic.

It has been suggest d that metals could ba coated -with a biocompatible plastic, ceramic or oxide to 25. overcome the corrosion problem. However, coatings tend to wear off and are susceptible to delaminating and separating from the metal substrate, exposing the metal to body fluids. - `~

30. Generally, it is the industry practice to passivate the implant metal alloys. However, -~
passivation produces only thin, amorphous, poorly attached protective oxide ~ilms which have not proved .' -: :

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totally effective in eliminating the formation of corrosion products in the body, particularly in situations where fretting occurs in the body.

5. As implant metals, titanium alloys offer advantages over stainless steels because of their lower susceptiblity to corrosion in the body coupled with their high strength and relatively low modulus o~
elasticity. Upon cooling, the currently used Ti-6A1-4V
10. alloy trans~orms from a ~-structure to an ~ plus B
structure at about 1000C. This transition can be shifted to a lower temperatuxe by the addition o~ one or more suitable ~-phase stabilizers such as . :: : .
molybdenum, zirconium, niobium, vanadium, tantalum, 15. cobalt, chromium, iron, manganese and nickel.~ -~".":~ ~".
Some efforts have been directed toward the development of alloys that eliminate harmful metals.
For example, US patent 4,040,129 to Steinemann et al is 20. directed to an alloy which includes titanium or zirconium as one component and, as a second component, any one or more of: nickel, tantalum, chromium, molybdenum or aluminum, but does not recognise or --suggest any advantages from having a relatively low 25. elastic modulus, or advantages or disadvantages - ~-associated with high temperature sinteri~g treatments (at about I250~C), commonly employed to attach porous ~-~
metal coatings into which bone can grow to stabilize non-cemented, press-fit devices into the skeletal 30. structure.

Although Steinemann provides that copper, cobalt, nickel, vanadium and tin should be excluded, apart from `~

'~ ';

:; 2 their presence as unavoidable impurities, the patent indicates that it is pe~missible to have any or all of chromium, molybdenum and aluminum~ which are all ~;
believed to have potential long-term adverse effects, 5. present in the alloy as long as their combined weight ~ -does not exceed 20% of the total weight of the alloy.

US patent 4,857,269 to Wang et al relates to a ~ -titanium alloy for a prosthetic implant said to have 10. high strength and a low modulus. The titanium alloy contains up to 24 wt% of at least one isomorphous beta stabilizer from the group molybdenum, tantalum, zirconium and niobium; up to 3 wt% of at least one eutectoid beta stabilizer from the group iron, 15. manganese, chromium, cobalt or nickel; ancl optionally up to 3 wt% of a metallic ~-stabilizer from the group aluminium and lanthanum~ Incidental impurities up to .. ..' - .q.::
0.05% carbon, 0.30% oxygen, 0.02% nitrogen, and up to 0.02% of the ~utectoid former hydrogen are also 20. included. Although there is eiome discussion of having an elastic modulus (eg. Young"s modulus) around 85 GPa, the only examples of a low modulus (66.9-77.9 GPa) all contain 11.5 wt% Mo which is a potentially toxic element and undesirable for optimizing ~ :~
25. biocompatibility.

Other currently used metal alloys have similar drawbacks. For example, the commonly used Ti-6A1-4V ~`
alloy, with appropriate heat treiatment, of~ers some 30. degree of biocompatibility but has an elastic modulus of about 120 GPa. Although this elastic modulus is lower than other alloys and accordingly offers better load transfer to the surrounding bone, this modulus is " ~ ~ ; j. ' ~ : j i . ' , ~ ': . ,. ~ ' , .'~ ~ . . j .' ' ' `: ' i; ' i i i ~ . i i ', ~ ! ' , ; i i ;

r still significantly greater than desired. Moreover, the alloy contains aluminium and also vanadium, which is now suspected to be a toxic or carcinogenic material ~-when present in sufficient quantity.
5. :. .
Commercially available PROTOSUL 100 (Sulzer Bros.
Ltd) is a Ti-6A1~7Nb alloy which intentionally avoids the potentially adverse e~fects of vanadium toxicity by substituting niobium. However, the alloy still -~
10. contains aluminium and has an elastic modulus of about 110 GPa (15.9 x 10 psi) in heat-treated condition, and with a tensile strength of about 1060 MPa. :~

With medical prostheses being implanted in 15. younger people and remaining in the human body for longer periods of time, there is a need for an implant material with requisite strength and flexibility requirements, which does not contain elements which are suspected as having long-term harmful effects on the 20. human body. Desirably, the implant material should have a hardened surface or coating that is resistant to microfretting wear and gross mechanical wear. ;~

Similarly, cardiovascular medical implants have 25. unique blood biocompatibility requirements to ensure that the device is not rejected (as in the case of natural tissue materials for heart valves and grafts for heart transplants) or that adverse thrombogenic ~clotting) or haemodynamic (blood flow) responses are 30. avoided.

Cardiovascular implants, ~uch as heart valves, can be fabricated from natural tissue. These f~ f 7 ~ ~ :

- 6 - ;

bioprostheses can be affected by gradual calcification leading to the ev~ntual stiffening and tearing of the implant. -;. Non bioprosthetic implants are fabricated from materials such as pyrolytic carbon-coated graphite, pyrolytic carbon-coated titanium, stainless steel, cobalt-chrome alloys, cobalt-nickel alloys, alumina coated with polypropylene and poly-4-fluoroethylene. `~
10.
For synthetic mechanifral cardiovascular devices, properties such as the surface finish, flow characteristics, surface structure, charge, wear, and mechanical integrity all play a role in the ultimate 15. success of the device. For example, typical materials used for balls and discs for heart valves include nylon, silicone, hollow titanium, TEFLON (Trade Mark), polyacetal, graphite, and pyrolytic carbon. Artificial hearts and ventricular as~ist devices are fabricated 20. from various combinations of stainless steel, cobalt alloy, titanium, Ti-6A1-4V alloy, carbon fibre reinforced composites, polyurethanes, BIOLON (Trade Mark, DuPont), XEMOT~ (Trade Mark, Sarns/3M~, DACRO~
(Trade Mark), poiysulfone, and other thermoplastics.
25. Pacers, defi~rillators, leads, and other similar cardiovascular implants are made of Ni-Co~Cr alloy, ~-Co-Cr-Mo alloy, titanium and Ti-6A1-4V alloy, stainless `~
steel, and various biocompatible polymers. Stents and vascular grafts are often made of DAC~ON (Trade Mark) 30. stainlf2ss steel or other polymers. Catheters and guide wires are constructed ~rom Co-Ni or stainless steel wire with surrounding polymer walls.

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One of the most significant problems encounteredin heart valves, artificial hearts, assist devic~s, pacers, leads, stents, and other cardiovascular implants is the formation of blood clots 5. (thrombogenesis~. Protein coatings are sometimes employed to reduce the risk of blood clot formation. ;
Heparin is also used as an anti-thromboyenic coating.

It has been found that stagnant flow reyions in ;~
10. the devices or non-optimal materials contribute to the ~ r ~ormation of blood clots. Thes~ stagnant regions can -be minimized by optimizing surface smoothness and minimizing abrupt changes in the size of the cross section through which ths blood flows or minimizing 15. either flow interference aspects. While materials selection for synthetic heart valves, and cardiovascular implants generally, is there~ore dictated by a requirement for blood compatibility to avoid the formation of blood c~lots (thrombus), 20. cardiovascular implants must also be designed to optimize blood flow and wear resistance.

Even beyond the limitations on materials imposed :::: ~ :~::
by the requirements of blood compatibility and 25. limitations to designs imposed by the need to optimize blood flow, there is a need for durable designs since ~ 3 it is highly desirable to avoid the risk of a second surgical procedure to implant cardiovascular devices.
Further, a catastrophic failure of an implanted device 30. will almost certainly result in the death of the patient.

The most popular current heart valve designs ~' ' : ~ .
~ ~, ) 7 7 ~

include the St. Jude medical tilting disc double cusp (bi-leaf) valve. This valve includes a circular ring-like pyrolytic carbon valve housing or frame and a flow control element which includes pyrolytic carbon 5. half-discs or leaves that pivot inside the housing to open and close the valve. The two leaves have a low profile and open to 85 from the horizontal axis. -Another popular heart valve is the Medtronic-Hall 10. Valve wherein the flow control element is a single tilting disc made of carbon coated wi'h pyrolytic carbon which pivots over a central strut inside a solid titanium ring-like housing. A third, less popular design, is the Omniscience valve which has a single 15. pyrolytic disc as a flow control element inside a titanium housing. Finally, the Starr-Edwards ball and cage valves have a silastic ball riding inside a cobalt-chrome alloy cage. The cage is affixed to one side of a ring-like body for attachment to the heart 20. tissue. More recent designs include trileaflet designs and concave bileaflet designs to improve blood flow.

From the point of view o~ durability, heart valves made of low-thrombo~enic pyrolyte car~on could 25. fail from disc or pivot joint wear or fracture related to uneven pyrolytic carbon coating, fracture of the ball cage, disc impingement, strut wear, disc wear, ~i ;
hinge failure,? and weld failure. A more recent heart valve, the Baruah Bileaflet is similar to the St~ Jude 30. design but opens to 80 and is made of zirconium metal.
The valve has worked well over its approximately two-year history with roughly 200 implants to date in India. This performance can be partly attributed to ~ ~ '; ';; ' .,.' . ' , . ` ' ~ ' ;

2 ~ ~ o r7 ~ ~ ~

9 ~ ~ . . . .

the lower elastic modulus of zirconium (about 90 GPa) and the resultant lower contact stress severity factor (Cc of about 0.28 x 10 7m) when the disc contacts the frame. In contrast, pyrolytic constructions produce 5. contact stress severity factors of about 0.54 x 10 7m. ;-~-~

Although zirconium has worked well to date and can reduce contact stress ~everity, zirconium metal is relatively soft and sensitive to fretting wear. This 10. is partly due to hard~ loosely attached, naturally~
present passive oxide surface films ~several nanometers -~
in thickness) which can initiate microabrasion and wear o~ the softer underlying metal. However, this naturally present zirconium oxide passive film is ~
15. thrombogenically compatible with blood ancl the design ~ -is acceptable from a haemodynamic standpoint.
Therefore, while the zirconium bileaflet valve appears to meet at least two of the major requirements for cardiac valve implants, namely blood compatibility and 20. d~sign for minimum stagnant f]Low regions, the use of soft zirconium metal leads to a relatively high rate of fretting wear and leads to the expectation that the -~
valve may be less durable than one produced from materials less susceptible to fretting wear. Titanium 25. and titanium alloys present a similar limitation, and Co-Cr-Mo, stainless steel, and Co-Ni alloys have much ;~
greater elastic modulus.

There exists a need for a metallic cardiac valve 30. impliant that is biocompatible, compatible with blood in that it does not induce blood clotting and dioes not form a calcified scale, that is designed to minimize stagnant flow areas where blood clotting can be ~: :,;~ ,',' $ ~

: :,:'. j,':,' ',`.i ', ':,.`.,`','"''.',,' " ' ''.`'-' '' ' .'';:" ' i~

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-- 1 o initiated, that has a low elastic modulus for lower contact stress severity factors to ensure resistance to wear from impact, and that has a surface that is also resistant to microabrasion thereby enhancing ;. durability.

Heart diseases, many of which cannot be cured by conventional surgery or drug therapy, continue to be a leading cause of death. For the seriously ill patient, 10. heart replacement is often one of the few viable options available. `~

Recently, research and development has been ~`
carried out into permanently implantable, electrically 15. driven, total artificial hearts (TAHs). The pumping mechanism of the TAHs would be implanted into the chest cavity of the patient and the device would be powered by a battery pack and a small transformer, worn by the patient, which transmits energy to the heart with no 20. physical connections through the skin.

~he development of TAHs posed several issues.
Firstly, it was necessary to duplicate the action of a -~
human heart, ensure long-term reliability and ~5. biocompatibility, while producing a device that fits into the chest cavity in terms of both its total volume and the orientation of its connections to natural ves~ls in the body. Aside from the purely mechanical, wear~ and power supply issues, it is also necessary 30 that the design and materials prevent infection and ~;
thromhosis. Blood is a non-Newtonian fluid and its properties, such as viscosity, change with oxygen content, kidney infection, and even the age of the 2 ~ 7 ~

patient. Further, plasma contains a suspension of fragile red blood cells which may be caught in artificial valves, or other mechanically stressful areas, thereby destroying these cells. It is therefore 5. necessary to develop a TAH that does not stress blood components, and to fabricate the pump from materials that are not only biocompatible, but also "blood compatible" in the sense of minimizing damage to blood components and minimizing the formation of blood clots. -~-1 0 . '' ~ ' Many of the above comments also apply to ventricular assist devices (VADs), one of which is being developed by the Novacor Division of Baxter Health Care Corp. In the use of a VAD, the patient's 15. hear~ remains in place while the VAD boosts the pumping pressure o~ the left ventricle o~ the heart.
Consequently, the VAD is an aie;sist device rather than a replacement. However, the VAO must be blood compatible for the same reasons as the total artifical heart.
20.
There exists a need for a material that is lightweight, readily formable into complex shapes, biocompatible, and blood and tissue compatible with a hard surface that is resistant to abrasive wear, 25. microfretting wear, and the corrosive effects of body fluids, for use in heart assist or replacement devices (including EMHs, VADs, and TAHs) to prolong the life of mechanical components while at the same time minimizing any deterious effect on blood components. -30.
According to the invention we provide a biocompatible medical implant of low modulus and high strength for implantation into a living body where it ;~ ?

7 ~

- 12 ~

is subject to corrosive effects of body fluids, said medical implant comprising~

a metalli~ alloy consisting essentially of:

(i) titanium;
(ii) from about 10 to 20 wt% niobium or from about 35 to about 50 wt% niobium; and ~ iii) optionally up to about 20 wt% zirconium;
1 0 .
wherein said medical implant has an elastic modulus less than about 90 GPa and the corrosive effects of body ~luids does not result in release of toxic or potentially toxic ions into the living body 15. after surgical implantation of the implant into said body.

The invention provides novel medical implants fabricated from hot worked, high strenqth, low modulus ~ -20. alloys of titanium, niobium and zirconium. The alloys are preferably free of toxic or potentially toxic compositions when used as an implant fabrication alloy.
More specifically, the invention alloy comprises titanium and niobium and optionally zirconium. To 25. achieve the lowest modulus, the titanium should ~`
preferably be alloyed with from about 10 wt% to about ~ ;
20 wt% niobium or from about 35 wt% to about 50 wt%
niobium. Zirconium is an optional compon~nt preferably ~-present in an amount from about 0 to about 20 wt%.
30. Most preferably the invention alloy comprises about 74 wt% titanium, about 13 wt% zirconium and about 13 wt%
niobium.

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7.3 By the term medical implant we mean, eg. kneejoint, consisting of a femoral component and a tibial base component; but particularly cardiovascular implants, that is device~ for implanting and use within S. the cardiovascular system; and implants for use in `~
conjunction with orthopaedic implants, eg. bone plates, ~
bone screwi~i, intramedullary rods and compression hip ::
screws. -~

10. The invention also provides cardiovascular `~
implants of a low modulus, biocompatible, hemocompatible, metallic alloy of titanium with niobium and optionally zirconium including heart valves, artificial hearts, ventricular assist devices, lS. defibrillators, pacers, electrical leads, sensors, grafts, stents, and catheter clevices. The invention also provides surface hardenecl versions of these davices produced by oxygen or nitrogen diffusion hardening to improve resistanc:e to cavitation, 20. microfretting wear, and impact-induced wear.

The inherently low modulus of Ti~Nb-Zr alloys, between about 6 to about 12 million psi depending on metallurgical treatment and composition, provide a more 25. flexible and forgiving construct for cardiovascular applications while improving contact stress levels, valve closure, and the ability of leaves in certain valve designs to self-align with blood flow and reduce thrombodynamic efects. ;~
30.
The invention provides components for use in mechanical heart replacement or assist devices, such as ~ :
external mechanical hearts ~EMHs), total artificial ~.

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hearts (TAHs), and ventricular assist devices (VADs), -that are lightweight, while also being resistant to corrosive body ~luids, mechanical wear, abrasive wear, and microfretting wear. Further, the components of 5. reduced risk of thrombogenesis (blood clotting).

The preferred low modulus titanium alloys of the invention for use as cardiovascular implants have the compositions: (i) titanium; about 10 wt% to about 20 -~
10. wt~ niobium; and optionally from about 0 wt% to about 20 wt% zirconium; and (ii) titanium; about 35 wt% to -about 50 wt% niobium; and optionally from about 0 wt%
to about 20 wt% zirconium. ~antalum can also be present as a substitute for Nb. These alloys are 15~ referred to herein as "~i-Nb-Zr alloys", even though tantalum may also be present.

The exclusion of alements besides titanium, zirconium, a~d niobium, or tantalum results in an alloy 20. which does not contain known toxins or carcinogens, or elements that are known or suspected of including diseases or adverse tissue response in the long term.

Without the presence of zirconium in the 25. composition, the ability of the Ti-Nb-Zr alloy to surface harden during oxygen or nitrogen diffusion hardening treatments is more limited. Therefore, pres~nce of zirconium i5 especially preferred when the -~
alloy implant must be diffusion hardened. Qther 30. non-toxic filler materials such as tantalum, which stabilize the ~phase of titanium alloy, but do not affect the low modulus, (the modulus of elasticity), ie. maintain it at less than about 85 GPa, could also -~ 2 ~ ~77~

- 15 ~

be added. The alloy may be substantially in the ~-phase and have a strength of greater than 620 MPa.

A porous coating, such as plasma-sprayed or ;. sintered titanium or titanium alloy (including Ti-Nb-Zr alloy) beads or wire mesh may also b~ added to the implant's surfaces to improve tissue attachment, such as the formation of an endothelial cell layer, preferred in artificial heart, ventricular assist 10. devices, grafts, and stent devices. Such coatings provide more favourable blood interaction and flow characteristics, and also tend to stabilize the implant with the body. Thus, such porous coatings may also be useful for connecting regions of these devices as well ~ -15. as for heart valves and grafts. Even though the ~ ;~
application of such porous coatings usually requires sintering at relatively high temperatures, the properties of the Ti-Nb-Zr alloy that might affect it~
usefulness as an implant material are not adversely~
20. affected.
: :
The invention alloy is ~trengthened by a hot working process wherein the alloy is heated to a temperature about its B-transus, or within about 100C `~
25. below its ~-transus, hot worXed, and then cooled rapidly, following which it is aged at temperatures below the B-transus. Preferably, this aging is carried out for about 2 to about 8 hours, most preferably about 6 hours at about 500C. The aging process may also 30. consist of a gradual ramp-up from room temperature, preferably in about 0.5 to 10 hours, during which preaging of the materi~l may occur, followed by iso~hermal aging at an appropriate temperature below ~ ;'` .'`'..'~
'"~-~: ~''' .~i .,.~.

7'~ 3 th~ ~-transus for from about 15 minutes to 20 hours, preferably about 6 hours. Note that by the time the alloy is removed from the furnace and the hot working operation performed, the temperature of the alloy may ;. have decreased signi~icantly; so the hot worXing operation may actually occur at a temperature significantly lower than the temperature to which the alloy is heated prior to hot working. The invention's hot worked, quenched and aged titanium alloys have a 10. low elastic modulus (about 60 to about 90 GPa) and have tensile strengths exceeding about 700 MPa, preferably exceeding about 800 MPa.

Further, the invention implants may ~e surface 15. hardened by any one of several processes used in the field of metallurgy but not necessarily known for use with medical implants. For example, there are processes in which the alloy is subjected to nitrogen or oxygen diffusion, internal oxidation, or nitrogen or 20. oxygen ion implantation. A description of these techniques may be found in our copending application, ;~
US Serial No.832,735, filed 7 February 1992, which is hereby incorporated by re~erence as if fully set forth.
Clearly, the alloy should be formed into the desired 25. 5hape for its intended use be~ore surface hardening or the benefits of surface hardness may be lost in any ~
subsequent shaping operations that affect or remove the ~;
surface of the alloy to a significant extent. Further, the shaped alloy may be worked for strength enhancement 30. before surface hardening.
,~ "
The most preferred hot worked, low modulus, high strength alloy for making medical implants contains ~: , ., ~,: : ~:
: . -, .

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- 17 ~

about 74 wt~ titanium, and about 13 wt~ each of zirconium and niobium. Other elements are not deliberately added, but may be present in trace amounts to the extent that they were present as unavoidable 5. impurities in the metals used to produce the alloy.
Other non-toxic filler materials such as tantalum, which could be used to stabilize the ~-phase, but not affect the low modulus (ie. maintain it less than about 90GPa), could also be added. The exclusion of elements 10. besides titanium, zirconium and niobium or tantalum results in an alloy which does not contain known toxins or carcinogens or elements that are known or suspected -~ ~
of including diseases or adverse tissue response in the ~ ~;
long term. Such an alloy is particularly useful in lS. medical implant applications.

The medical implant according to the invention may comprise at least a partial outer surface -protective coating selected from the group consisting 20. of the oxides, nitrides, carbides and carbonitrides of elements of the metal alloy. It may further comprise a protective coating of amorphous diamond-liXe carbon on at least a portion of an outer surface of the implant.

25. The metallic alloy may be internally oxidised or ~-nitrided beneath outer surfaces of the implant to produce a hardened medical implant. ~ -;

We also provide a heart valve prosthesis for ~-30. implantation in living bodOv tissue of a patient, the heart valve having enhanced hemocompatibility,~ ;
comprising~

:

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(a) a valve body having an aperture through which blood is able to flow when the heart valve is implanted in a patient, the valve body fabricated from a metal alloy comprising:
;. (i) titanium;
(ii) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (iii~ optionally up to about 20 wt% zirconium;

10. (b) a flow control element able to move relative to the valve body to close the aperture in the valve body thereby blocking blood flow through the aperture; ~
and : ~:

15. (c) means, attached to the valve body, for restraining said flow control component to close proximity to the aperture in the valve body.

We also provide a ventr:icular assist device 20. including components with surfaces in contact with blood when the device is implanted in a patient, the improvement comprising:

said components fabricated from a metal alloy 25. comprising:

(i) titanium; :~
(ii) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and 30. ~iii) optionally up to about 20 wt~ zirconium.

A total artificial heart device is provided for implantation into a chest cavity of a patient, the . ~
,,, ~

:
: :~

rJ~ 7 ~3 device including components with surfaces subject to mechanical wear and microfretting wear when in use in the patient, the improvement comprising:

5. (i) titanium;
(ii) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and ~:
(iii) optionally up to about 20 wt% zirconium.

10. In a total artificial heart device for implantation into a chest cavity of a patient, the device including components pr~senting surface~ that are in contact with blood when in use in the patient, ~ `
the imp:rovement wherein the components are fabricated 15. from an alloy comprising:

(i) titanium;
(ii) ~rom about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and 20. (iii) optionally up to about 20 wt% zirconium.

We also provide a flexible guide wire for~:
insertion into a living body to perform surgical operations, the guide wire comprising:
25. :
(a) an elongate, flexible guide wire body having a distal end, for insertion into a patient and into a catheter, and a proximal end for controlling the guide wire; -~
30.
(b) an elongate guide wire core-disposed internally along the longitudinal axis of the elongate body~ said core comprising a metal alloy comprising:

~ 21~ nrl7~

(i) titanium;
(ii) from about 10 to about 20 wt% niobium orfrom about 35 to about 50 wt% niobium; and (iii) optionally up to about 20 wt% zirconium. ~
i . , - . .
There is also provided an expandable stent for ~ ~:
supporting a blood, urinary, or gastrointeskinal vessel from collapsing inward, the stent comprising~

10. (a) a radially outwardly expandable substantially cylindrical stent body of a metal alloy ;: ~-comprlslng~

(i) titanium;
15. (ii) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and ' (iii) optionally up to about 20 wt% zirconium.

the stent body in its unexpected state sized for 20. ease of insertion into a vesse!l and in expanded state : :
sized for propping open a vess;el, ths stent body having a bore therethrough for recei~ing a means for expanding ~:
the body radially outwardly. ~-~

25. A biocompatible lead or sensor for conducting ~ :
electrical signals to or from an organ in a living ~ ~ .
body, the lead comprising~

an elongate flexible body having distal and 30. proximal ends, the flexible body comprising~
~ .
~ a) an electrically conductive core of a metal alloy comprising~

~` 21:~'7~l3 (i) titanium;
(ii) from about lo to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (iii) optionally up to about 20 wt% zirconium.
5.
A low modulus, biocompatible, percutanaous implant is provided that penetrates the skin of a living body and thereby protrude~ from the body, the implant comprising~
1 0 .
(a) a low modulus metallic implant body fabricated from a metal alloy comprising:

(i) titanium;
15. ii) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (iii) optionally up to about 20 wt% zirconium.

said implant body havin~ a first end for 20. insertion into said patient and a portion with a second end for extending outside of s,aid patient.

There is provided an external mechanical heart including components with surfaces subject to z5. mechanical wear and microfretting wear, the improvement comprising:

components of a metal alloy comprising:

(a) titanium;
(b) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (c) optionally up to about 20 wt% 2irconium.

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``` ~3:~10779 In an external mechanical heart including ~ ;
components that contact blood when said heart is used to supply blood to a patient, the improvement comprising the components fabricated from a metal alloy `~
;. comprising~

(a) titanium;
(b) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and 10. (c) optionally up to about 20 wt~ ~irconium.

A vascular graft of enhanced durability, crush-resistance, low thrombogenicity, and hemocompatibility, said graft comprising:
15.
(a~ titanium;
(b) from about 10 to about 20 wt% niobium or -from about 35 to about 50 wt% niobium; and (c) optionally up to about 20 wt% zirconium.
20.
said tubular body having a bore therethrough, said tubular body sized to replace a section of removed blood vessel.

25. The cardiovascular devices may comprise a metal alloy comprising from about 0.5 to about 20 wt%
zirconium. The outer sur~aces of the devices may be hardened by a process selected from the group consistin~ of oxygen diffusion hardening, nitrogen 30. hardening, physical vapour deposition, and chemical vapour deposition. The devices may further comprise a coating overlayed over outer surfac~s comprising a medi ament. In addition it may comprise wear-resistant ~ ;~

: ~ ~ e~ L ~ 7 ~

surfaces produced by a procPss selected from the group ~ ;
consisting of boronation and silver doping.

The invention's hot worked, low modulus, high `~
5. strength titanium implants are produced by heating to ahove the B-transus temperature (or within the range including those temperatures below and within less than about 100C of the B-transus temperature); hot working the implant; rapidly cooling the hot worked alloy to 10. about room temperature; then reheating and aging at t~mperatures below the B-transus, in the range of 350-550C, prPferably about 500C, for a time ~-sufficient to provide an implant of adequate strength.

15. In a particularly preferred aging process, the implant is preaged by gradually heating the quenched implant for a period o~ time up to about 350-550C.
Thereafter, the preaged implant is i~othermally aged at this temperature for a time sufficient to develop 20. 5trength and hardness characteristics required.

The invention implants have a low modulus of elasticity of less than about 90 GPa. This is a significant improvement oYer Ti-6A1-4V which has a 25. modulus of elasticity of a~out: 120 GPa.

In certain applications it may still be desirable ~-to coat the implant's surface with wear-resistant - .
coatings such as amorphous diamond-like carbon ~-30. coatings, ~irconium dioxide coatings, titanium nitrides, carbides, or the like for protection against potenti~l micro fretting wear such as might occur on -~
the bearing surfaces of implant prostheses.
''...' ~'.

J ~ 7 ~

A porous coating, such as a bead, powder, or wiremesh coating may be applied to implants of many types or a variety of applications fabricated from the inventive alloy. Such coatings are often usaful to 5. provide interstitial spaces for bone or tissue ingrowth i~
into the implant, which t~nds to stabilize the implant in the skeletal structure.
-While implants fabricated from the invention hot 10. worked alloy possess hiyh strength, the usefulness of i~
these prostAeses is not limited to load-bearing applications. Because of its corrosion resistance ~ ;
non-toxicity and relatively low modulus of elasticity, the alloy can be used to fabricate many types of 15. medical implants including, but not limited to, knee joints, cheek bones, tooth implants, skull plates, fracture plates, intramedullary rods, staples, bone screws, spinal implants, pelvic plates, and other implants, cardiovascular implants such as synthetic 20. heart valves, ventricular assist devices, total artificial hearts, stents, grafts, pacers, pacemaker leads and other electrical lealds and sensors, defibxillators~ guide wires and catheters, and ~j ~
percutaneous devices. ~ -25.
Figure 1 shows a simplified representation of a -ball valve, like the Starr-Edwards Valve.

Figure 2 is a simplified representation of a disc 30 valve. -;

Figure 3 is a simplified representation of a tilting di6c, single cusp valve like the Medtronic-~all ~"'''. ~

7 7 ~

valve.

Figure 4 is a simplified representation of a tilting disc, double cusp or bileaflet valve o~ the St.
;. Jude or Baruah type.

Figure 5 is a schematic diagram of the Penn State/Sarnst3M design total artificial heart~

10. Figure 6 is a schematic di~gram of the University of Utah total artificial heart.

Figure 7 is a schematic diagram of the Cleveland Clinic/Nimbus Inc total artificial heart.
15.
Figure 8A is a schematic diagram, in cross section, of the Texas Heart Institute/Abiomed total artifical heart.

20. Figure 8B is a side view, in cross section, of the Texas Heart Institute/Abiomed total artifical heart ~ -of Figure 8A.

Figures 9A and B are schematic diagrams of end 25. and front views, respectively, of a ventricular assist device.

Figure lO is a schematic diagram of a vascular -graft of woven metallic wire composition.
30.
Figure llA is a schematic diagram showing a balloon expandable stent positioned within a segment of a blood vessel to be propped open.

' '~ ~', ,'~

!7 7 ~3 ' Figur2 llB is a schematic diagram showing aballoon expanding stent into position within a blood vessel.

;. Figure llC is a schematic diagram of a stent expanded in a blood vessel.

Figure 12A is a schematic diagram of the components of a defibrillator, showing power source, 10. lead wire, and polymeric patch with coiled electxode.

Figure 12B is a cross section oP the lead wire of Figure 12A.

15. Figure 13A is schematic diagram, in partial cross section, of the distal end of a guide wire.

Figure 13B is a cross section of the guide wire with coating thickness exaggerated.
20.
Figure 13C shows a catheter containing coiled wire and polymer wall.

Figures 14A-C are schematic diagrams of prior art 25. pacemaker leads with polyurethane covering.

Figure 14D is a schematic of an embodiment of the invention Ti-Nb-Zr pacemaker leads.

30. Figure 15 is a schematic diagram of a modular knee joint prosthesis.

Figure 16A is a schematic diagram of a side view '~,'';' ~'' :~ , -~ ~"''' ~`- 2~773 `

of a typical intramedullary rod used in orthopaedic application~.

Figure 16B is a view of 16A, rotated by 90 .
~. .
Figure 17A is a schematic diagram of a typical bone screw.

Figure 17B is an end view of 17A.
10 .
Figure 18 is a schematic diagram showing a typical screw for fixing bone plates.

Figure 19 is a schematic diagra~ of a compression lS. hip screw.

Figure 20A is a side view of a typical bone plate, in partial cross-section, for securing to the hip.
20.
Figure 20B is the plate of 20A rotated by 90 to show 6 holes for bone screws to affix the plate in the :~:
body and a topmost hole for receiving a compression hip screw, like that of Figura 19.
25.
Figure 20C is a cross-sectional view of 20B taken at 9C-9C.

The inventive alloy implants may be produced by 30. combining, as commercially pure components, titanium, niobium and optionally zirconium in the appropriate proportions, heating the alloy to above its ~-transus (or within the range of temperatures below and within ~`

-'` 211~ 9 about 100C of the B-transus), hot working, rapidly cooling to about room temperature, and aging the alloy at temperatures below the ~-transus for a su~ficient length of time to allow strength development. It is 5. essential that cooling be carried out rapidly, as by ~uenching with water. Conventional convective air cooling is not sufficiently rapid to produce the high strength, low modulus alloy of the invention after aging.
10. ~ ' Heart Valves In its simplest form, a synthetic cardiac valve ~;~
includes a valve body for affixing the valve to the 15. body tissue and through which blood flows, and a flow control element for allowing or blocking off blood flow. For instance, Figure 4 shows a typical bileaflet valve having a valve body that includes a ring-like housing 400 with an inner ring 402 that has two flanges 20. 404 each containing two ~lots ~or receiving hinges attached to leaflets. The flow control element of this valve comprises two leaflets 405, in the approximate shape of half discs, with hing~ elements attached at ~ ~r-diametrically opposite ends. These hinge elements fit 25. within apertures or slots in the flanges 404 of the inner ring 402 and are able to rotate through less than 180C in these apertures. Thus, in operation, the flow control elements are in the position shown in Figure 4 with th~ valve open with blood flowing from top to 30. bottom. When blood flow reverses and flows ~rom bottom to top, the bileaflets 405 pivot about their hinges to close the apertures in the ring-like valve body. ~;
Conse~uently, there is a significant amount of movement -7 7 ;3 about the hinge elements and slots where microfrettingwear might be initiated. Furthermore, the bileaflet half disc flow control elements 405 may impinge upon the inner ring 402 of the valve body, thereby leading 5. to cavitation or impact-induced wear.

The invantion provides heart valves of various designs, exemplified in Figure 1-4, each comprising parts subject to impact and wear that are fabricated 10. from Ti-Nb-Zr alloy. More recent designs (not shown in :-the Figures) include concave bileaflet and trileaflet designs which are intended to improve blood flow. The heart valves are preferably subjected to a hardening process, such as oxygen or nitrogen diffusion hardening 15. to produce a harder Ti-Nb-Zr surface that is resistant .
to microfretting wear at hinge points and impact wear at those locations where a flow control element impacts the valve body. Consequently, the invention valves have a longer cycle li~e than the currently used St.
. Jude, Omniscience, Starn-Edwards, Medtronic-Hall, or ~Y
Baruah valves. Indeed, as mentioned ~efore, the Baruah :~
valve is currently fabricated of zirconium or zirconium alloys and would therefore be subject to relatively rapid wear because of the relative softness of . zirconium and its alloys. The use of Ti-Nb-Zr alloy compositions also provides a low thrombus, blood-compatible surface fa~ourable in reducing the incidence of blood clots.

30~ Artificial Hearts/Ventr ular_Assist Devices :~

Figure 5 is illustrative o~ the Penn State design which incorporates a stainless steel roller screw 10 f ~

positioned between two flexible diaphragm blood pumps (right side pump 12 is shown, the other is within shell 8). A high speed, low torque brushless DC motor causes the roller screw 10 to turn thereby moving the roller 5. screw shafts 16 linearly back and forth. To each end of the guide shaft 16 is ~ittached a pushPr plate ~8.
When the pusher plate 18 in the right side pump moves bacXward, blood is drawn into the pump space 14. At the same time, the pusher plate in the corresponding 10. left side pump moves towards the left pushing blood out of its pump space. In this type o~ pump, the pusher plates 1~ act against a flexible membrane 12, which is in contact with blood, and which can be in~lated on the pump suction stroke and deflated on pump discharge 15. stroke- The pusher plates are driven by the roller screw 10 aind, thus, according to the invention, six ~
revolutions of roller screw 10 are required for a full r stroke with a motor speed of about 3,000 RPM. While planetary rollers are inserted between a roller screw ~ ;
20. nut aind roller screw 10 to give rolling, not sliding contact and to spread mechanical load over many contact points, and while the entire roller screw system is improved by the use of the invention components. Thus, ~-roller screw 10 and the roller screw nut with which it 25. is in rolling contact arP both fabricated of Ti-Nb-Zr alloy and the surfaces are hardened or coated with a hard, tightly adherent coating. Further, ~he surfaces of the pump nozzles 2 shown as 4 and 6, connecting elements 20 and conduits 22, into and from which blood 30. is continuously being pumped, is fabricated from Ti-Nb-Zr to reduce adverse reaction with blood tissue on those sur~aces presented to blood components to minimize the potential for the ~ormation of thrombus and blood clots.

Figure 6 is a schematic cross sectional diagramof the u~iversity of ut~h electrohydraulic heart. This 5. heart includes shells 50 and 52 with a pump motor 34 interposed between them. In this type of' heart, a motor 20 is used to pressurize silicone oil in regions Z2 and 24 on the undersides of flexi~le diaphragms 26 and 28, respectively, to move blood in and out of the 10. chambers 30 and 32 above the flexible diaphragms. For example, when motor 34 pressurizes silicone oil into chamber 22, then flPxible diaphragm 26 expands upward :~
and outwardly to push blood flow out of chamber 30 in direction 36. At the same time, silicon oil flows out .
15. Of chamber 24 towards chamber 22 thereby allowing ~:
flexible diaphragm 2B to assume a natural position, :
shown in Figure 2, and drawing blood into chamber 32 as :~
shown from direction 38. Upon reversal of the direction of the bidirectional pump 34, the opposite 20. effects are achieved.

Since the University of Utah pump is of a bidirectional de~ign, and typically operates at speeds between 10,000 - 13,000 RPM in the high pressure 25. direction and 5,000 8,000 RPM in th~ rQverse, moving :~:
components of the pump are subject to microf'xetting and mechanical wear. Therefore, the invention components or the bidirectional axial flow pump 34 used in the University of Utah TAH design are fabricated from Ti-Nb-~r alloy coated hardened or with a wear resistant, hard coating that is tightly adherent, to reduce wear of the high speed components. Further, surfaces 40, 41, 42, 43, 44 and 45 are in direct 2 1 1 0 7 7 ~

contact with blood and are made of Ti-Nb-Zr alloy to improv~ blood compatibility and reduce the potential for thxombus and blood clotting. Thus, the shells of the heart 50 and 52 are also fabricated of Ti-Nb-Zr ;. alloy to reduce thrombogenesis.

Figure 7 is a schematic cross sectional . :~
illustxation of the Cleveland Clinic TA~ which utilizes a motor to turn a gear pump 56 which provides hydraulic 10. pressure at about 100 psi to caus~ reciprocal movement of actuator~ 58 which in turn drive pusher-plates 60 that act on flexible diaphragms 62 to pump the bloodO `~
The actuators 58 operate slidingly within guide sleeve 64 so that wear on contact surfaces between actuator 15. and sleeve may be expectedO Further, the TAH has a flow reversing valve 64 with machine elements, such as bearing surfaces, subject to wear when the TAH is in us~. Thus, TAH elements that are subject to wear and that may be advantageously fabricated of Ti-Nb-Zr :
20. alloys that are then surface hardened and/or coated with a hard, wear resistant, tightly adherent coating, ~ :.
include the guide sleeve 64, the actuators 58, the ~:
pump's gear elem~nts and shaft and the rotary valve 64.
Moreov~r, to reduce the risk of erosion damage to the 25. pump from cavitation, the pump housing 72 may likewise ~ :
be fabricated of Ti-Nb-Zr alloys. Finally, internal i ~
surfaces 66, 68 of the heart housing 70 are in direct :~;
blood contact. Thus it is desirable to fabricate housing 70 from Ti-Nb-Zr to reduce the risk of 30. thrombogenesis.

Figures 8A and B are schematic diagrams of the T~xas Heart Institute/Abiomed T~H design which utilizes ',';."" '..'.,.,.,..:',,': "".-.,""'``'''.'."'' '"'"'"'''~'' ' " ~'' "'"' `" :' h ~L ~ ~) 7, ~

a d.c. motor to drive a miniature centrifugal pump 80 that rotates at about 6,000 - 8,000 RPM. This pump 80 pressurizes hydraulic fluid alterna~ely into chambers 82 and ~4 separated by septum 86 and enclosed ~y 5. flexible diaphragms 88 and 90 respectively. As fluid is pumped into chamber 82, diaphragm 8B expands into heart space 92 forcing blood from this spaceO At the same time, fluid is pumped from chamber 84 causing diaphragm 90 to relax and expanding heart space 94, 10. drawing blood into the TAH. The hydraulic flow is ~ ~ ~
reversed by a two-position 4-way rotating valve 100 of ~:
radial configuration for compactness. Rotary valve 100 rotates within sleeves 102 and 104 at high speed so that contacting surfaces between the valve 100 and 15. these s:Leeve5are subject to wear. Further, rotary valve 100 rotates against seals 106 and wear may be expected at the contacting surfaces of the seals and the valve 100.

. Several components of the Texas Heart Institute/
Abiomed TAH may be fabricated according to the invention. Thus, high speed components of the centrifugal pump 80 subject to wear may be fabricated from Ti~Nb-Zr alloy and then surface hardened and/or 25. coated with a tightly adherent, hard, wear resiistant coating. Further, the rotary valve 100 itself and the sur~aces of sleeves 102, 104 and seals 106 may be fabricated fro~ Ti-Nb-Zr alloy then sur~ace hardened or coated with an adherent, wear-resistant coating.
30. Finally, the inner surfaces of the TAH 108, 110 may be fabricated from Ti-Nb-Zr alloys to improve blood compatibility and r~duce the poten-tial for thrombus and blood clotting.

2 1 L V 7 ~ 3 -- 3~ --The Novacor designed VAD illustrated in Figures 9A and ~ have a solenoid mechanism 120 which sends energy through beam-springs 122, 124 that extend to the back of pump pusher plates 126~ 12~. The energy stored 5. in the springs translates into lin~ar motion of the plates which Pxerts force on the flexible blood sac 130. The blood sac 130 consists of a butyl rubber layer sandwiched between two layers of polyurethane Biomer. The blood sac 130 is supported within a 10. cylindrical aluminium ring 132 that acts as a pump housing. The blood inflow 133 and outflow 134 ports are positioned tang~ntially on opposite sides of the -~
housing to ensure straight-through blood flow~ The ports are formed of an epoxy-impregnated Kevlar fabric 15. shell that is integrated into the housing. The ports also encapsulate trileaflet inlet and outlet valves made from bovine pericardium tissue. When implanted into the body, fittings for attaching inflow and outflow valves to vascular conduits are bonded to a 20. pump bulkhead, not shown, which also provides the framework for an encapsulatin~ shell around the pump.
This encapsulating shell also has provision for mounting the solenoid energy converter. The solenoid energy converter consists of two solenoid mechanisms, 25. two lightweight titanium beam-springs~ and an aluminium support structure. ~ll of these metallic components would come into contact with blood components and body issue. Therefore, the invention proposes that the titanium beam-springs be replaced with beam-springs of 30. Ti~Nb-Zr alloy. Further, the aluminium support structure would likewise be replaced with a Ti-Nb-Zr alloy support structure that may optionally be hardened and/or coated with a hard coating.

2 ~ 7 ~ ~

Novacor has identified, in desiqning the solenoid, that "the challenge was coming up with -~
something that would run for lOO million cycles a year, without requiring maintenance". O'Connor, Lee, ~-~
5. "Novacor's VAD: How to Mend a Broken Heart", Mechan.
Engr'g pp. 53-55 (Nov 1991). The i~vention components fabricated from Ti-Nb-Zr alloys then hardened or coated with hard, wear resistant coatings provide surfaces that are hard, microfretting wear resistant, lO. biocompatibls and blood compatible so that they would ~ -meet this goal. To further reduce friction and wear of wear sur~aces of implant devices, a thin boron or silver surface layer can be applied as an overlay on the previously diffusion hardened Ti-Nb-Zr surface.
15.
External mechanical hearts (EMHs) are used as a bridge to transplant. These heaxts include the Jarvik-7 pneumatic heart and the more recent left-ventricular assist device, the Heartmate developed 20. by Thermocardio Systems. In the Heartmate system, two tubes, one carrying air and the other electrical wire, pass from outside the body to an implanted blood pump.
The pump is implanted in the abdomen and removes blood from the natural heart's left ventricle. This blood 25. enters and exits the pump through 25 millimeter input and output valves made from chemically processed bovine tissue. T~e blood flows from the output valve through a dacron-wrapped polyurethane tube to ths aorta. An electric motor mounted in the Heartmate's lower chamber 30. actuates a ~lat-plate piston, which is bonded to a flexible polyurethane diaphragm. When the motor goes through one revolution, it turns a cam assembly that compresses the diaphragm, which pushes blood through ;

21 ~f~ ~73 the output valve. The operation of the pump is controlled by a microprocessor located in a shoulder ~i bag which adjusts the heartbeat rate by changing the motor' 5 current. According to the invention, the 5. moving parts of the heartmate pump may be replaced with components ~abricated Prom Ti-Nb-Zr alloys then hardened or coated with a hard coating to reduce mechanical wear, friction, and microfretting wear.
Furthermore, those metallic components that come into 10. contact with blood components, may also be replaced with Ti-Nb-Zr alloy components similarly coated to -~
improve blood compatibility and reduce the risk of clot formation.

1~. The gravest problem in the use of the pneumatic Jarvik-7 heart has been identified as the formation of blood clots. O'Connor, Lee, "Engineering a Replacement for the Human Heart", Mechan. Engr'g pp. 36-43 (Jul 1991). In 1990, the FDA withdrew the Jarvik system 20. from clinical trials due to concerns over quality control during manufacture. The University of Utah made modiPications to the design of the Jarvik heart to develop a new system called the "Utah 100" which has elliptical pump housings, as opposed to the spherical 25. housings of the J~rvik-70 Further, the Utah 100 has redesigned junctions for joining the diaphragms within the ventricles to the housing. These changes are said to have resulted in an about 70~ reduction in blood clot formation relative to the Jarvik-7 de~ign.
30. However, according to the invention, yet further reduction in blood clot formation may be obtained by fabricating moving parts and those metallic surfaces that contact blood components from Ti-Nb-Zr alloys and 2 ~ 7 '~ ~

-- 37 -- :

then hardening and~or coating these components with hard, wear resistant coating to increase blood compatibility and thrombus resistance, and to reduce abrasive wear, and reduce microfretting wear.
5.
Guide Wires and Cathete~s Figures 13A and B show, in partial cross section, the distal end of a guide wire fabricated according to 10. the invention. The guide wire 145 has a core 140 of Ti-Nb-Zr alloy with a surface hardened coating 142 to reduce friction which may be further coated with a material that is bio- and hemocompatible and of low ~riction when in contact with the catheter wall or body 15. tissue. The flexible catheter 146 through which the guide wire moves is also fabricated according to the invention and includes a coil 147 of low modulus titanium alloy which is generally encased by a polymer sheath 148 as shown schematically in Figure 13C. The 20. guide wire in this cass is equipped with a cutting tip 144, preferably also made of Ti-Nb~Zr alloy with a dif~usion hardened surface optionally with a hard ceramic or lubricating coating. Since the guide wire is fabricated of metal, it is highly visible under 25. x-rays, providing excellent radiopacity. Boron or silver surface layers may also be deposited on the diffusion hardened surfaces to further reduce friction and wear.

30. Pacemakers and Electrical Siq~ Carryin~ Leads~Sensors Pacemaker and other electronic leads are manufactured by several corporations, including ~:

2 1 1 0 7 ~ ~

Medtronic, which produces a range of pacemaker leaddesigns.

One of these designs is shown in schematic form S. in Figures 14A and B. The pacemaker lead body 150 has a centrally disposed metallic conductor 152 typically made of cobalt-nickel alloy, such as MP35N (Trade Mark). This conductor 152 is usually made up of several strands of wire, each having a diameter of 10. a~out 0.15-0.20 mm. The conductor 152 is covered by an insulative, protective polymer sheath 153 so that the elongate body 150 of the pacemaker lead has an overall diameter ranging from about 2.2 to about 3 mm. The pacemaker has a first end 154 with an electrode 158 for 15. connecting to a pulse generator and a serond end 156 with an electrode 157 for contacting heart muscle. An alternative embodiment is shown in Figure 14C. As supplied, thes~ two ends are covered with protective polyurethane caps which can be! removed ~or installation 20. of the pacemaker. In order to prevent electrical interference with the conductor 152, a polymeric insulative sleeve 153 is disposed over the entire pacemaker lead body 150, with the exception of th~
exposed el~ctrodes 157 for contacting heart muscle and 25. the contact electrode 158 for engaging with the pulse generator that hou~es the electronics and power pack for the pacemaker. As explained b~fore, the organic polymeric sheath compositions, typically polyurethane, can ~lowly degenerate in the body causing problems, not 30. only due to potential deterioration of electrical insulation and interf~rence with electrical signals but also because of potentially toxic products of degradation.

/ ,i~ 7 ~ i~

The invention provides, as shown in Figure 14C, a pacemaker wherein the conductor 15~ is fabricated from a Ti-Nb-Zr alloy that is coat~d with a tightly adherent, low friction, bio- and hem~compatibl2 5. coating, with the exception of the electrode for contacting heart muscle 157, and the electrode 158 at the other end of the lead for engaging the pulse generatorO The coatings can be formed by in situ oxidation or nitriding of the Ti-Nb-Zr to produce an 10. electrically insulative surface layer of from about O.1 to about 3 microns in thickness, preferably less than about 0.5 microns in thicknesæ. This process can be carried out at the same time the material is age-hardened. Alternatively, an insulative inert 15. ceramic coating can be appliecl by conventional CVD or PVD m~thods either on the original Ti-Nb-Zr alloy surface or onto the diffusion hardened Ti-Nb-Zr surface. For these overlay coatings, the thickness can be as great as 20 microns. The overlay coatings 20. include ceramic metal oxides, metal nitrides, metal carbides, amorphous diamond like carbon, as detailed above. The electrical signal conductor 152 can comprise either a single wire or multiple wires.
Exposed Ti-Nb-Zr metallic ends of the wire or wires are ~5. preferably connected directly to a pulse generator thereby avoiding the necessity for a weld or crimp to attach an electrode to the conductor which may result in local galvanic corrosion or physically we~kened regions. Further, since the coatings provide ~ natural 30. protective insulative surface, the use of a coiled construct could be avoided by using only a preferred ;
single-strand, non-coiled low modulus Ti-Nb-Zr metallic wire construct for th~ conductor 152. This will also .3 40 ~

elimina e the need for stiff guide wire. Finally, the overall diameter of the pacemaker lead body 150 could be reduc~d considerably from the range of about 2.3 - 3 mm for current commercially available leads to about ;. 0.2 - 1 mm. Optionally, the leads of the invention may be covered with a polymieric sheath.

Stents 10. Figure llA shows a schematic of an expandable stent ~60, in non-expanded state, positioned on thiP
distal end of a balloon expandable segment 162 of a guide wire 164. The stent is fabricated from Ti-Nb-Zr alloy and is designed so that it can be collapsed over lS. a balloon segment of a balloon catheter. When the stent is in position, within segment of a tubular conduit 165 in the body, a blood vessel for example, to be propped open, the balloon 162 is expanded thereby expanding the stent 160 radially outward up to the 20. blood vessel wall 166 so that means for gripping soft tissue, such as barbs (not shown), on the outer surface of the stent 160 engage and grip blood vessel tissue to anchor the stent 160 in po~ition as shown in Figure llB. The balloon 162 is then collapsed and removed 25. leaving ths stent in place as shown in Figure llC. In this way, the blood vessel is permanently propped open.
Urinary, gastrointestinal, and other stent applications are also provided using Ti-Nb-Zr alloy.

30. Grafts ~ igure 10 is a representative skeitch of a side view of a substantially tubular vascular graft 170 .

~ 3 41 ~

. sizPd to graft onto a blood vessel and made of woven low modulus Ti-Nb-Zr wires 172. While the graft shown is made of woven wires of Ti-Nb-Zr, th~ graft can also be fabricated from a cylindrical tubing of this alloy.
5. The graft can be fabricated from Ti-Nb-Zr alloy in the lower modulus cold worked condition, or in the slightly higher modulus aged condition with optional surface hardening. Additionally, protein, antibiotic, anti-thrombic, and other surface treatments may be ¦ 10. employed to further improve the biocompatibility and clinical performance.

nefibrillators 15. Figures 12A and B show a defibrillator including a ~lexible silicone polymeric patch 300 with a coil of conductive wire 320 (typically titanium, stainl~ss steel, or cobalt-nickel-chromium) on the side of the silicone patch 300 that will contact muscle tissue.
20. When in place in the body, the lead wire 320 that carries power to the coil 340 extends out of the body (through the skin~ and is electrically connected to a pow~r sour~e contained in a protective container 360.
According to the invention, the lead wire 320 is 25. fabricated with an electrically conductive core 350 of Ti-Nb-Zr alloy and is coated with an adherent ele~trically insulative coating 2B0, such as metal ~ :~
oxides, carbides, or nitrides or with amorphous diamond-like carbon as shown in exaggerated detail 30. Figure l~B. This Goating electrically insulates the lead wire from electrical contact with surrounding body ~ .
tissue while also protecting the me~allic core from corrosion and attack by body fluids, as described :

previously, for the pacemaker lead. Elimination of t~polymer coating results in the elimination of pot~ntially toxic products of gradual degradation of the polym~r and also th~ consequent shorting the system 5. when the insulative coating is breached.

The Hardened Surfaces The oxygen or nitrogen diffusion hardened surface 10. of the alloy implants may be highly polished to a mirror finish to further improv~ blood flow characteristics. Further, the oxide- or nitride-coated surfaces may be coated with substances that enhance biocompatibility and performance. For example, a 15. coating of phosphatidyl choline, heparin, or other proteins to reduce platelet adhe~ion to the surfaces of the implant, or the use of antibiotic coatings to minimize the potential for infection. Borona~ed or ;`~
~ilver-doped hardened surface layers on the implant 20. reduces friction and wear bet~een conta~ting parts of heart valves, prosthetic arti~Eicial hearts, ~entricular assist devices, and other contacting parts in the invention cardiovascular implants. Additionally, amorphous diamond-like carbon, pyrolytic carbon, or 25. other hard ceramic surface layers can also be coated onto the diffusion hardened surface to optimize other friction and w~ar aspects. The preferred diffusion hardened surface layer described in this application provides a hard, well attached layer to which these 30. additional hard caatings with respect to hardnesis.
Other, conv~ntional methods of oxygen surface hardening are ~lso u~eful. Nitriding of the substrate leads to a hardened nitride surface layer. M~thods of nikridation 2~ )779 ~3 - ;

known in the art may be used tc achieve a hard nitride layer.

Regaxdless of how a Ti-Nb-Zr alloy implant's 5. surface is hardened, the friction and wear (tribiological) aspects of the surface can be further improved by employing the use of silver doping or boronation techniques. Ion-beam-assisted deposition of silver films onto ceramic surfaces can improve 10. tribiological b~haviour. The deposition of up to ahout 3 microns thick silver films can be performed at room temperature in a vacuum chamber equipped w:ith an electron-beam hard silver evaporation source. A
mixture of argon and oxygen gas is fed through the ion 15. source to create an ion flux. One set of acceptable silver cleposition parameters consists of an acceleration voltage of 1 kev with an ion current density of 25 microamps per cm2. The silver film can be completely deposited by thi~s ion bombardment or 20. ~ormed partially via bombardmemt while the remaining thickness is achieved by vacuum evaporation. Ion bombardment improves the attac]hment of the silver film to the Ti-Nb-Zr alloy substratle. Similar deposition of silver ~ilms on existing metal cardiovascular implants 25. may also be performed to improve tribiological behaviour, a~ well as antibacterial response~

An alternative method to further improve the tribiological behaviQur of Ti-Nb-Zr alloy surfaces of 30. cardiovascular implants is to apply boronation treatments to these surfaces such as commercial available boride vapour deposition, boron ion implantation or sputter deposition using standard ion - 2 3 ~ 7 3 implantation and evaporation methods, or form a boron-type coating spontaneously in air. Boric Acid ~H3BO3) surface films provide a self replenishing solid lubricant which can further reduce the friction and 5. wPar of the ceramic substrate. These film~ form from the reaction of the B2O3 surface (deposited by various coventional methodsj on the metal surface with water in the body to form lubricous boric acid. Conventional methods that can be used to deposit either a baron ~B), ~
10. H3BO3, or B203 surface lay r on the cardiovascular :::
implant surface include vacuum evaporation twith or without ion bombardment) or simple oven curinq of a thin layer over the implant surface. The self-lubricating mechanism of H3BO3 is governed by its 15. unique layered, triclinic crystal structure which allows sheets of atoms to easily slide over each other during articulation, thus minimzing substrate wear and friction.

~o Additionally, surfaces (metal or coated) of all the cardiovascular and medica:L impl~nts discussed may optionally be coated with agents to further improve biological response. These agents include anticoagulants, proteins, ant:imicrobial agents, 25. anti~iotics, and the like medicaments.

The titanium alloys are also useful in the manufacture of medical implants, and possess the characteristics of high strength, low modulus of 30. elisticity, corrosion resistance to ~ody fluids and tissue, and are free from any potentially toxic elements. Thus, the alloys are especially useful in the fabrication of bone plates (Figures 20A,8,C), intramedullary rods (Figures 16A,B), compression hip screws (Figure 19), spinal implants, modular knee joints tFigure 15), and the like. Typic~l modular knee joints as shown in Figure 15 include a femoral :
5. component 240 and a tibial component 250. The femoral component includes condyles 242 which provide ::
articulating surfaces and pegs 244 for affixing to the femur. The tibial component 250 includes a tibial bas~
252 with a peg 254 for mounting the base onto the :
10. tibia. A tibial plat~orm 256 is mounted atop the tibial base 252 and is supplied with grooves 258 that cooperate with the condyle~ 242. The tibial platform :~
256 is ~requently made of an organic polymer (such as ultra-high molecular weight polyethylene) but the 15. tibial base 252 and femoral component 240 are ~:
~abricatPd of metal. The invention provides tibial bases and ~emoral components of the above-described titanium-niobium-zirconium alloys.

20. The preferred titanium alloys for medical implants for use other than in the cardiac system include: (1) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium, and t2) optionally up to 20 wt% zirconium.
25.
The most preferred inventive alloy for medical use outside the cardiac system contains titanium as the major component comprising about 7~ wt% Qf the alloy in combination with about 13 wt% of zirconium and 13 wt%
30. of niobium. -~

While tantalum may be substituted for niobium to sta?bilize the ~-phase titanium, niobium is the ~ ' '' '"~ ~'.`.

L `L ~ 3 - 46 ~

preferred component due to its effect of lowering the elastic modulus of the alloy whan present in certain specific proportions. Other elements are not deliberately added to the alloy but may be present in 5. such quantities that occur as impurities in the commerically pure titanium, zirconium, niobium or tantalum used to prepare the alloy and such contaminants as may arise from the melting (alloying) process. Filler materials, such as non~toxic tantalum, 10. could also be added to reduce the ~-transus ~stabilze B) and improve strength a~ long as the relatively low modulus of elasticity tless than about 90 GPa~ of the base alloy is not significantly affected.

15. Based upon the foregoing, it is apparent that the titanium proportion of certain embodiments of the invention alloy could be less than 50 wt%.
Nevertheless, these alloys are, for purposes o the speci~ication and claims, refe!rred to as "titanium 20. alloys". For example, a titanium alloy may contain 20 wt% zirconium and 45 wt% niobi.um with only 35 wt%
titanium.

While the as-cast or powder metallurgically 25. prepared alloy can be used as an implant material or for other applications, it can optionally be mechanically hot worked at 600-950~Co The hot working process may include such operations as extrusion, hammer f~rging, bending, press forging, upsetting, hot 30. rolling, swagging, and the like. After the final hot working step, the alloy should be cooled rapdily, as for instance, by water quenching. Slower rates of cooling, as by air convection cooling, are not ~:~

.'' t~

recommended and are not effective in producing the highstrength, low modulus alloy suitable for use a~ a medical implant. After cooling, it can then be reheated, preferably gradually over a period from about S. 0.5 to 10 hours, more preferably 1.5 to 5 hours to a maximum temperature of about 700C, preferably about 500C. Then, the implant material is maintained at this temperature for from about 0.25 to 20 hours, preferably for from about 2 to about 8 hours, more 10. preferably for about 6 hours. This proc~ss, by a pheno-menon called precipitation strengthening, is responsible for the high strength o~ the alloy in the hot worked, quenched and aged condition described above.
1~.
In the specification and claims, the term "high - :
strength" refers to a tensile strength above about 700 MPa, preferably above about 800 MPa.

20. The term "low modulus" as used in the specification and claims refer to a Young's modulus below about 90 GPa. : :

In titanium alloys, the niobium (or tantalum, if ~ r 25. this element is added) acts to stabilize the ~-phase since it is a B-isomorphous phase stabilizer. This :~:
results in a lower B-phase transus temperature and ~:
improved hot wor~ability. ~
, - :, 30. Niobium, in particular, when present in preferred quantities of from about ~ to about 10 atomic percent (most preferably about 8 atomic perGent) or in an alternative pr~ferred range of from about 22 to ~2 :::

'i ~^ " i~,',:, ~.' f. L ~ ~ 7 7 ~

atomic percent, produces a low modulus composition when alloyed with titanium. Deviation from these ranges of niobium concentration tends to increase the elastic modulus. In weight percent terms, these preferred 5. compositional ranges of niobium in the titanium-zirconium alloy translate to about 10 -to about Z0 wt%
and about 35 to about 50 wt%.

Titanium alloys containing about 13 wt% niobium 10. correspond to those having about 8 atomic percent niobium. Thus, the Ti~13Nb-13Zr alloy is believed to id~ntify an optimal low modulus, titanium alloy compositlon.

15. As previously mentioned, tantalum may be substituted for niobium to stabilize the ~-phase, but niobium is preferred due to its effect in reducing the elastic modulus. Substitution with zirconium can improve strength.
20.
Whereas the niobium prop,ortion is critical to i~
obtain the desired low modulus property, the zirconium proportion is not as critical. It is desirable to maintain th~ proportion of zirconium at less than about 25. 20 wt% but higher proportions are also useful. -~:' Zirconium, it is believed, is capable of stabilizing both ~ and ~-phase titanium alloy, but acts by being in solution in the alloy as a B-stabillzer by slowing the transformation process in the inventive alloy. It is further believed that the larger ionic radius of zirconium (35% larger than that of titanium) helps to disrupt ionic bonding forces in the alloy r~J !`t ~ 49 --resulting in some reduction in the modulus of elasticity.

In order to ~ffect the translation to the B-phase ;. (which is not essential to produ~e the high strength, low modulus alloy implants of the invention), the alloy may be treated by heating to above the ~-transus temperature, egO to about 875C, for about 20 minutes.
Lower temperatures above the B-transus may also be 10. used. The ~-phase may also be induced by heating to ::
higher temperatures for shorter periods of time. The critical factor for transition to the ~-phase is heating to at least about the ~-transition temperature, which is about 728C for Ti-13Zr-13Nb, for a period of 15. time su:fficient to obtain a substantial conversion of the titanium alloy to the ~-phase prior to cooling to :
room temperature. Conversion of the alloy to the B-phase and cooling may be e~fected be~ore, durinq or ~.
after shaping Eor implantation and sintering of a .,-20. porous metal coating, whichever is most convenient.

It should be noted that heating to the above the ~-transus and hot working at such elevated temperature while converting most or all of the alloy to the 25. B-phase, is not essential to obtain the desired high strength and low modulus. Indeed, the alloy may be ~:
heated to a temperature as low as about 100C below the B-transus, and hot worked at this lower temperature to ~ -achieve high stxength and low modulus, after rapid 30. quenching and aging, without complete transformation to the B-phase. The region immediately below the ~-transus temperature is called the ~ ~ B region. Hot working may be performed after heating to this ~

~ "~""',''';~ `

2 f 1 ~ ~J ~ ~

region and, possibly, even temperatures below this region, ie. temperatures as low as about 100C below the B-transus.

5. The ~-transus temperature of the most preferred Ti-13Nb-13Zr alloy is about 728 C. The alloy may be heated to above the B-transus, eg. about 800C, for forging. Other intermediate temperatures may also be used, but at temperatures lower than about 600C
10. forging may be difficult because o~ the poorer ~ormability of the alloy at these low temperatures.

The machining, casting or forging of the alloy into the desired implant shape may be carried out by 15. any of the conventional methods used for titanium alloys. Further, implants could be pressed from the powdered alloy under conditions of heat and pressure in pre-forms in the shape of the desired implant.
Conventional sintering and hot isostatic pressure 2Q. treatments can be applied.

While the alloy provides a non-toxic prosthesis material, it may yet be desirable for other reasons, such as micro-fratting against bone or polyethylene 25. bearing surfaces, to coat th~ metal surface. In this event, the surface may be coated with an amorphous diamond-like carbon coating or ceramic-like coating such as titanium nitride or titanium carbide, or the oxide, nitride or carbide of zirconium, using chemical 30. or plasma vapor deposition techniques to provide a hard, impervious, smooth surface coating.

Alternatively, a coating may be formed in situ on the shapPd alloy by exposure to air, oxygen, and/ornitrogen at elevated temperatures to oxidize or nitride the surface of the alloy to a deslred depth. Typically these coatings, resulting from the diffusion of oxygen ;. or nitrogen into the metal surface, are up to about loO~ thick or greater. These in situ coatings are tightly adherent and more wear resistant than the metallic alloy sur~ace. Coatings are therefore -especially useful if the alloy is subjected to 10. conditions of wear, such as, for instance, in the case ~i~
of bearing surfaces of knee or hip pro~theses.

Methods for providing hard, low-friction, impervious, biocompatible amorphous diamond like carbon 15. coatings are known in the art and are disclosed in, for example, EP0 patent application 302 717 A1 to Ion Tech and Chemical Abstract 43655P, Volume 101 describing Japan Kokai 59/851 to Sumitomo Electric, all of which are incorporated by reference herein as though fully 20. set forth.

Further, the metal alloys may be hardened by interstitial ion implantation wherein the metal surface is bombarded with the ions of oxygen or nitrogen, and . the like. The metal retains a metallic-appearinq surface but the surface is hardened to a depth of about 0.1~. The metals may also be surface hardened by internal oxidation, as described in our copending US
Serial No.832,735, filed 7 February 19920 30.
Implants fabricated from the inventive alloy may be supplied with a porous bead, powder, or wire coating oE titanium alloy of the same or different compocition , .. , .. .. .. ...... , ,...... j.. . .

,. 2~ 7'~9 - 52 ~

including pure titanium to allow stabiliæation of the implant in the skeletal structure of the patient after implantation, by bone ingrowth into the porous structure. Such porous structures are sometimPs 5. attached to the implant surface by sintering. This involves heating the implant to above about 1250C.
The mechanical properties of titanium alloys can change significantly due to substantial grain growth and other metallurgical ~actors arising from the sintering 10. process. Thus, after sintering to attach the porous coating, it is preferred that the Ti-13Zr-13Nb implant be reheated to about 875C (ox above the ~-transus) for 20-40 minutes then quenched before being aged at about 500C for about 6 hours to restore mechanical 15. properties. If quenched adequately from the sintering temperature, it may be possible to go directly to the aging process.
~:
The ~ollowing examples are intended to illustrate 20. the invention as described above and claimed hereafter and are not intended to limit the scope o~ the invention in any way. The aging temperature used in ~he examples is determined to ~e acceptable, although not necessarily optimal 25.
Example 1 An alloy including, by weight~ 74% titanium, 13% ~-niobium and 13% zirconium, was hot rolled at a 30 temperature in the range 825-875C to 14mm thick plate.
The plate was cooled to room temperature then reheated to 875C where it was maintained for 20 minute~ and then water quenched to room temperature. The B-transus ~ 2 ~ 77~

for this alloy was about 728C as compared to about 1000C for Ti-6Al-V. The mechanical properties of the heat-treated, quenched Ti-Zr-Nb alloy, which has an .:~
acicular trans~ormed B-structure, are shown in Table I. -~

~AB~E I
Mechanical Properties of Ti-13Zr-13Nb As water Quench~d from above B-Transus Temperature Tensile Strength 710 MPa 10. Yield Strength 476 MPa Elongation 26% :~
Reduction in Area 70% i:~
Young's Modulus 62 GPa ::
Rockwell C Hardness 18-19 1 5 .
Example 2 The heat-treated, quenched Ti-Zr-Nb alloy of Example 1 was aged by heating at 500C for 6 hours.
20~ The mechanical properties of this aged alloy are shown in Table II.

T~Bh~
Mechanical Properties of Quenched 25. Ti-13Zr 13Nb A~ed 500 C for Six Hours Tensile Srength 917 MPa Yield Strength 796 MPa :~
Elongaticn 13%
Reduction in Area 42%
30. Young's Modulu~ 76.6 GPa Rockwell C Hardness About 29 ~lQ77~

Example 3 Samples of the alloy of Example 1 were sintered at about 1250C to attach a porou~ titanium bead 5. coating of the type shown in Figure 1. The bead-coated alloy samples were then reheated to 875c and maintained at this temperature for 40 minutes before being water~quen~hed~ A group of six samples were aged at 500C for 6 hours and the mechanical properties of 10. aged and non-aged samples (three each~ were tested and shown in Table III.

T~BL2 IIX
M~echanical Properties of Ti-13Zr-13Nb Alloy :~:
15. Following Bead Sintering, Reheating to 875C, and Water Quenched As-Quenched (Av~L _ Aqed ~500C Six Hours) Tensile Strength 664 MPa 900 MPa Yield Strength 465 MPa 795 MPa 20. Elongation 20% 4%
Reduction Area 46% g%
Young's Modulus 61.8 GF~a 74.7 ~Pa Note that the sintering treatment can 25. significantly alter the mechanical properties, particularly ductilityO Thus, an alloy acceptable for a particular application in unsintered form may not ~ :
necessarily be effective in that application following a high-temperature sinteriny treatment often used to 30. attach a porous~titanium coating. To minimize these effects, lower temperature diffusion bonding methods can be used in which a temperature near the B-transus may be effective. Alternatively, pre-sintered porous 7 ~

metal pads can be tack-welded to the implant. Yet another alternative is to apply the psrous coating by a plasma-spraying m~thod which does not expose the bulk of the material to high temperature.
5.
Example 4 A comparison of the elastic modulus, tensile strength and yield strength of the Ti-13Zr-13Nb 10. invention alloy with those of known alloys, composites and cortical bone, are summarized in Figures 2 and 3.
A1203 and ZrO2 refer to ceramics while C/PEEK refers to carbon reinforced polyetheretherketone composite and C/PS refers to a carbon reinforced polysulfone 1~. composite. All the mechanical property data of Figures 2 and 3 were obtained from literature sources except ~or the data pertaining to the! invention alloy which were measured using standard ASTM t~nsile testing techniques. It i5 significant that the Ti-13Zr-13Nb 20. invention alloy ha~ an ela~tic modulus similar to carbon fibre reinforced composites and closer to that of bone than the other metals (Figure 2~ while at the same time possessing a strengt.h comparable to or better than other metals (Figure 3)0 25.
xample 5 A sample of Ti 18Zr-~Nb was sintered to attach a porous metal coatingO Thereafter, the sintered alloy 30. was reheatad to 875C, ie. above the B-transus, and water quenched. The properties of the as-quenched alloy are shown in Table IV. The sample was then aged a~ 450C for 3 hours and tested. These results are -: ~; 2i-~ ~77~

also shown in Table IV.
:
As compared to the Ti-13Zr-13Nb alloy of Example 3, this alloy'~ modulus of elasticity is not as low but 5. is still lower than that of Ti-6~1-4V. Further, the Ti-18Zr-6N~ alloy has a relatively low B-transus, about 760C, compared to that of Ti-6A1-4V which is about 1000C.

10. TABL~ IV
Mechanical Properties of Ti-18Zr-6Nb Following A
High Temperature Sintering Treatment, Reheating to ~75C, and Water Quenchlnq and Aclina As-Quenched Aqed 450C, 3 Hrs 15. Tensile Strength 807 MPa 876 MPa Yield Strength 659 MPa 733 MPa Elongation 8% 8%
Reduction in Area 26% 28%
Elastic Modulus 85.2 GPa 86~8 GPa 20.
Note that because of the less than optimum niobium content, the elastic modulus is not as low as th2 pr~vious example. Thus, proper selection of niobium content is important for optimiæing the low 25. elasti~ modulus. However, the presence of zirconium helps to keep the elastic modulus at an acceptably low :
level (less than about 90 GPa).

Example 6 30. :
The effect of aging conditions on Ti-13Zr-13Nb ~.
and Ti-18Zr-6Nb was investigated. Separate sample of each alloy were air-cooled or water-quenched from above :~

7 1 ~

~ 57 -., .,~

the ~i-transus, aged at 500, 450, 400 and 350C for up to 6 hours then air cooled. The results are recorded in Figure 4.

5. Example 7 Forgings of Ti-13Nb-13Zr were prepared at temperatures of 800 and 680C. These forgings were :
either water quenched or air cooled from the ~orging 10. temperature and their mechanical properties were :
determined:

TABLE V
Forqin~ Tem~ C 800 800 680 680 800*
15. Quench medium Air Water Air Water Water Ultimate Tensile 789 765 852762 691 Strength, MPa Yield Strength, MPa 592 518 700517 44~
Elongation, %21 23 18 27 28 20. Reduction in Area, %72 60 /2 73 72 Young's Modulus, GPa84 66 85 72 60 Hardness, Rockwell C20 26 27 25 * Net-shape forged ~5.
Forgings of 19mm minimum diameter Ti-13Nb-13Zr, which had been air cooled after forging, were heat treated by aging at 350 to 500C for from 1.5 to 6 hours. The mechanical properties of these forgings 30. were as follows:

!.~:: 'r,~'," ~. .. ' j~ ,~ . ' ~ .~ ~ " ~ ' i ':. ~'~

2 ~ 7 7 ~

~a Forqing Temp, C 800 800 _ 800 ~00 :~
Aging Temp, C 500 450350 350 Aging Time, hrs 6 1.5 6 1.5 ;. Ultimate Tensile 775810 g72 882 Strength, MPa Yield Strength, MPa 651625 641 639 Elongation, % 23 ~0 18 17 Reduction in Area, % 74 70 60 59 10. Young's Modulus, GPa 86 84 88 87 The strength of these alloys is not very high, even after aging, since they had been air cooled (slow cooling as opposed to rapid cooling of water quenching) 15. after f~rging.

Other Ti-13Nb-13Zr forgings, also of minimum l9mm diamet~r, were water quenched after forging, then heat treated. The mechanical properties of thase forgings 20. are recorded in Table VI.

25.

'~

30. ~:

'7'~ 9 ~ ~e: ~ ~ ~3_i 8~ 2~ s~_ a~ ~ I
1~ - 3 '` ~ ~ ~ ~ I

t--~ -- N N = ` O 1~ ~ I ~ I ~ ' I

O N O ~~ _i _ t~i ~ t~

o c~i u~ ~ioti ~ ~ ~ ._ ~ri ai ai ,_ _ _ _ _ __ _ . _ .

OEi _i O Nii i _ i O ~ _ ~ ~
O i O O ¢ ~ i `J It~ ~ ~c~

_ _ _ 0 _ ~ _ _ ai ~i a t.~ ~ ~ ~- ~ ~ 1 33~

l .. ci E f.~ f~ ai C ~: ui ~ _._ C r C~ C~ Li~ S~ ~i V '~i C-D8ii ~ ~ ; i a 77~

The alloy shows significant further improvementin strength while retaining a Young's modulus in the range 74-90 GPa~ Hardness is also significantly greater than for the non-heat treated alloy as well as ;. the reheated, quenched, and aged alloy of Example 2.

Example 8 Several tests were performed to determine the 10. effect on the phy5ical properties of Ti-13Zr-13Nb when forged at different temperatures, aged at different temperatures for periods of time, and quenched under different conditions.

15. Water quenched l9mm minimum diameter samples from hip stems (proximal, mid, or distal sections) were heat treated at low temperatures for short times and their properties measured and recorded in Table VII.

.
20.

25. ~:

30. ~

~'' . ~, ' ~ -L L ~ r~ ~ ~3 ~ 61 ~
~o 3n~ ~ ~ ~ 0 _~0 ~I :~

o ~ c~ c~ _~ _ ~ O co _~ r~ o 1~ - ~- -- ~
cl ~ o ~ ~ ro o ~ o ~ :~: c~r~ ~ o o~ I~
l _ __ _ _ __ ~ o~ o~ eo ~ ~
~_ ~ _ _ :-o ~ ~r r~

~1 c~ r~ O c~J S ~ 3~ C~l ".~ ~ .
~ O ~'~ O~ ~D ~ CI~ _ ,, : _ _ ._ _ _ __.. _ .~ ~U ~
a o~ , ~ ~ ,~ c ~ c .
~= _ : . ~__ 1~ J
O ~ O O 1 N Ul ~ ~ ~

. ~. o o o o o I q e ~ ~
I _ I o gr~
_ ___ ___ ~=; __ _ ~_ .. __ _ _ ~ ~
~ ~ ~ :~ ~ ~ C~ q,0~ U~ ~
l _ . h ~ _ _ 1: ~ O ~ ~ ~ ~ ~ : ~
I ~D. e~ e c~ D _~ ~ h~ X
I 0~ r~ ~ _~ 1~ U U 0~ _ U~ I I I I I I
L c c _ ~ _ _ _ ~ 3 _ ~ ~ ~ --~ ~ :

'~

77~

Forgings of 14mm minimum diameter were w~erquenched aEter the final Porging step only, then aged at 500C for 1, 2 or 3 hours. Their properties were:

~AB~B VIII
.
Forgings Temp, C 800 ~00 800 Aging Temp, C 500 500 500 Aging Time, hrs 1 2 3 Ultimate Tensile 953 960 1017 10. Strength, MPa Yield Strength, MPa 802 811 898 Elongation, % 14 11 12 Reduction in Area, % 50 44 42 Young's Modulus, GPa 73 73 ~2 15.
The forgings show an increase in both yield s-trength and tensile str~nyth with time of aging.

Example 9 20.
A 12.5 inch long by 14 i,nch diameter segment was cut from an ingot of Ti-13Zr-13Nb which had been ~ ;
produced by arc melting. The plate was press forged to a 3 inch thickness at 1000C. The press forged plate 25. was then air annealed at 1100C for 15-30 minutes be~ore being hot rolled at 900C to a 1.35 inch ~:
penultimate thickness. The hot rolled plate was then r*heated to 900C, hot rolled to 1.04 inch final thirkness, and water quenched be~ore being blasted and 30. pickledO The mechanical properties of the plate were as follows~

-` 2 Ll~'~79 TABLB I~
_late Ultimate Tensile Strengith (MPa) 786+0 Yield Strength (MPa)539~12 5. Elongation (%) 21+0 Reduction in Area (~) 51+5 Young's Modulus (GPa) 74+3 Hardness, Rockwell C24.S+2.0 .
10. The hot worked and water quenched plate was then subjected to aging cycles consisting of a gradual heating up step and an isothermal aging step. The heating up appeared to produce a "preaging e~fect" on - -the material which enhanced subsequent aging response 15. and produced a higher strength. The aging cycle included heating up the plate over a period of 1.5 to 5 hours up to 500C (preaging), followed by 6 hours of ..
aging at 500C. The resultant mechanical properties of the plate are shown in Table IXB:
20.
ABL~ I~B

25.

30.

S 7 .~ ~

- ~4 From comparing Tables IXA and IXB, it is apparent that strength and hardness have increased significantly due to the aging process.

5. Example 10 A 54 inch long by 14 inch diameter bar was cut from an ingot of Ti-13Zr-13Nb produced by arc melting.
The bar was rotary forged to a 5.4 inch diameter at 10. 875C, then air annealed at 1050C f~r 3 hours.
Thereafter, the bar was rotary forged at 800~C to 2.5 inch diameter, and rotary forged at 750C to 1.2 inch penultimate diameter. The swagged bar was then reheated to 925C before being rotary swagged to a 1 15. inch final diameter. The bar was water quenched, blasted and pickled~ and then centrelPss ground.

Th2 resultant mechanical properties of this Ti-13Zr-13Nb bar were as follows:
20.
T~LE ~
Bar Ultimate Ten~ile Strength (MPa~ 722+9 Yield Strength ~MPa~463+10 25. Elongation ~) 26+0 Rsduction in Area (%~66+2 Young's Modulus (GPa~79+9 Hardnoss, Rockwell C2~+2.7 30. Th~ hot worked and water quenched bar was then subjected to aging cycles including a gradual heating up step and an .isoth2rmal aging step. Once again, the heating up appeared to produce a preaging ef~ect on the ,:~

~,3 ~ ~ ~ }~ r~

bar, which enhanced subsequent aging response and produced a higher strength material. During aging, the bar was first gradually heated, over a p~riod of ~.5 to about 5 hours, up to 500C. This was followed by 6 5. hours of isothermal aging at 500c. The resultant mechanical properties of the bar were as follows-T~B~E ~ -Ramp-up Time to 509C thr). 2.5 10 .
Ultimate Tensile Strength (MPa) 1008+6 Yield Strength (MPa)881+13 Elongation (~) 10+3 Reduction in Area (A%)32+14 15. Young's Modulus (GPa)82.4+2.3 Hardness, Rockwell C31.5+0.6 As can be seen from a comparison of Table XA and Table XB, the strength and hardness of the bar ;
20. increased significantly as a result o~ the aging and preaging processes.

The invention has been described with reference to its preferred embodiments. From this description, a 25. person of ordinary skill in the art may appreciate ~, changes that could be made in the invention which do not depart from the scope and spirit of the invention as described above and claimed hereafter.

30.

Claims

1. A biocompatible medical implant of low modulus and high strength for implantation into a living body where it is subject to corrosive effects of body fluids, said medical implant comprising:

a metallic alloy consisting essentially of:

(i) titanium;
(ii) from about 10 to 20 wt% niobium or from about 25 to about 50 wt% niobium; and (iii) optionally up to about 20 wt% zirconium;

wherein said medical implant has an elastic modulus less than about 90 GPa and the corrosive effects of body fluids does not result in release of toxic or potentially toxic ions into the living body after surgical implantation of the implant into said body.

2. The medical implant according to claim 1 which is a cardiovascular implant.

3. The medical implant according to claims 1 or 2 consisting essentially of:

(i) titanium;
(ii) from about 10 to 20 wt% niobium; and (iii) up to about 20 wt% zirconium.

4. The medical implant according to claims 1 or 2 consisting essentially of:

(i) titanium;

(ii) from about 35 to 50 wt% niobium; and (iii) up to about 20 wt% zirconium.

5. The medical implant of claim 1 wherein the implant is selected from the group consisting of bone plates, bone screws, intramedullary rods, and compression hip screws.

6. The medical implant of claim 1 wherein the metallic alloy comprises tantalum in an amount sufficient to stabilise the .beta.-phase without significantly affecting the modulus of elasticity of the implant.

7. The medical implant of claim 1 wherein the alloy is substantially in the .beta.-phase and has a strength greater than about 620 MPa.

8. The medical implant of claim 1, wherein the alloy consists essentially of: 74 wt.% titanium, 13 wt.%
niobium, and 13 wt.% zirconium.

9. The medical implant of claim 8 wherein the alloy is substantially in the .beta.-phase and has a strength greater than about 620 MPa.

10. The medical implant of claim 1 further comprising at least a partial outer surface protective coating selected from the group consisting of the oxides, nitrides, carbides and carbonitrides of elements of the metal alloy.

11. The medical implant of claim 1 further comprising a protective coating of amorphous diamond-like carbon on at least a portion of an outer surface of the implant.

12. The medical implant of claim 1 wherein the metallic alloy is internally oxidised or nitrided beneath outer surfaces of the implant to produce a hardened medical implant.

13. The medical implant of claim 1 wherein the implant is selected from the components of a modular knee joint consisting of a femoral component and a tibial base component.

14. The medical implant of claim 13 wherein the metallic alloy is internally oxidised or nitrided beneath outer surfaces of the implant to produce a hardened medical implant.

15. The medical implant of claim 13 wherein the alloy is substantially in the .beta.-phase and has a strength greater than about 620 MPa.

16. The medical implant of claim 13 wherein the alloy consists essentially of: 74 wt% titanium, 13 wt%
niobium, and 13 wt% zirconium.

17. The medical implant of claim 13 further comprising a protective coating of amorphous diamond-like carbon on at least a portion of an outer surface of the implant.

18. The medical implant of claim 13 further comprising at least a partial outer surface protective coating selected from the group consisting of the oxides, nitrides, carbides and carbonitrides of elements of the metal alloy.

19. The medical implant according to claim 2 comprising a heart valve prosthesis for implantation in living body tissue of a patient, the heart valve having enhanced hemocompatibility, comprising:

(a) a valve body having an aperture through which blood is able to flow when the heart valve is implanted in a patient, the valve body fabricated from a metal alloy comprising:

(i) titanium;
(ii) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (iii) optionally up to about 20 wt% zirconium;

(b) a flow control element able to move relative to the valve body to close the aperture in the valve body thereby blocking blood flow through the aperture;
and (c) means, attached to the valve body, for restraining said flow control component to close proximity to the aperture in the valve body.

20. The heart valve prosthesis of claim 19, wherein the metal alloy comprises from about 0.5 to about 20 wt.% zirconium.

21. The heart valve prosthesis of claim 20, wherein outer surfaces of the valve body are hardened by a process selected from the group consisting of oxygen diffusion hardening, nitrogen hardening, physical vapour deposition, and chemical vapour deposition.

22. The heart valve prosthesis of claim 19 further comprising a coating overlayed over outer surfaces comprising a medicament.

23. The heart valve prosthesis of claim 19 further comprising wear-resistant surfaces produced by a process selected from the group consisting of boronation and silver doping.

24. The medical implant according to claim 2 comprising a ventricular assist device including components with surfaces subject to mechanical wear and microfretting wear, the improvement comprising:

components fabricated from a metal alloy comprising:

(a) titanium;
(b) from about 10 to about 20 wt% niobium or from about 25 to 50 wt% niobium; and (c) optionally up to about 20 wt% zirconium.

25. The ventricular assist device of claim 24, wherein the metal alloy comprises from about 0.5 to about 20 wt% zirconium.

26. The ventricular assist device of claim 25, wherein the surfaces subject to mechanical wear and microfretting wear are hardened by a process selected from the group consisting of oxygen diffusion hardening, nitrogen hardening, physical vapour deposition, and chemical vapour deposition.

27. The ventrical assist device of claim 24, wherein the surfaces are coated with a medicament.

28. The ventricular assist device of claim 24 further comprising a coating on the surfaces subject to mechanical and microfretting wear, the coating applied by a process selected from the group consisting of silver doping and boronation.

29. The medical implant according to claim 2 comprising a ventricular assist device including components with surfaces in contact with blood when the device is implanted in a patient, the improvement comprising:

said components fabricated from a metal alloy comprising:

(a) titanium;
(b) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (c) optionally up to about 20 wt% zirconium.

30. The device of claim 29, wherein the metal alloy comprises from about 0.5 to about 20 wt% zirconium.

31. The device of claim 30, wherein the surfaces in contact with blood are hardened by a process selected from the group consisting of oxygen diffusion hardening, nitrogen hardening, physical vapour deposition, and chemical vapour deposition.

32. The device of claim 29, wherein the surfaces in contact with blood are coated with a composition comprising a medicament.

33. The medical implant according to claim 2 comprising a total artificial heart device for implantation into a chest cavity of a patient, the device including components with surfaces subject to mechanical wear and microfretting wear when in use in the patient, the improvement comprising:

components fabricated from a metal alloy comprising:

(a) titanium;
(b) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (c) optionally up to about 20 wt% zirconium.

34. The device of claim 33, wherein the metal alloy comprises from about 0.5 to about 20 wt% zirconium.

35. The device of claim 34, wherein the surfaces subject to mechanical wear and microfretting wear are hardened by a process selected from the group consisting of oxygen diffusion hardening, nitrogen hardening, physical vapour deposition, and chemical vapour deposition.

36. The device of claim 33, wherein the surfaces subject to mechanical wear and microfretting wear are hardened by coating with a process selected from the group consisting of silver doping and boronation.

37. The medical implant according to claim 2 comprising a total artificial heart device for implantation into a chest cavity of a patient, the device including components presenting surfaces that are in contact with blood when in use in the patient, the improvement wherein the components are fabricated from an alloy comprising:

(a) titanium (b) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (c) optionally up to about 20 wt% zirconium.

38. The device of claim 37, wherein the metal alloy comprises from about 0.5 to about 20 wt% zirconium.

39. The device of claim 38, wherein outer surfaces of the components are hardened by a process selected from the group consisting of oxygen diffusion hardening, nitrogen hardening, physical vapour deposition, and chemical vapour deposition.

40. The device of claim 37, wherein the surfaces in contact with blood are coated overlayed over outer surfaces comprising medicament.

41. The medical implant according to claim 2 comprising a flexible guide wire for insertion into a living body to perform surgical operations, the guide wire comprising:

(a) an elongate, flexible guide wire body having a distal end, for insertion into a patient and into a catheter, and a proximal end for controlling the guide wire;

(b) an elongate guide wire core disposed internally along the longitudinal axis of the elongate body, said core comprising a metal alloy comprising:

(i) titanium;
(ii) from about 10 to about 20 wt% niobium or from about 35 to 50 wt% niobium; and (iii) optionally up to about 20 wt% zirconium.

42. The guide wire of claim 41 further comprising a hardened cutting edge on the distal end of the elongate body.

43. The wire of claim 41, wherein the metal alloy comprises from about 0.5 to about 20 wt% zirconium.

44. The wire of claim 43, wherein the core has a hardened outer surface produced by a process selected from the group consisting of oxygen diffusion hardening, nitrogen hardening, physical vapour deposition, and chemical vapour deposition.

45. The wire of claim 41, where surfaces of the wire that come into contact with body tissue are coated with a composition comprising a medicament.

46. The wire of claim 41, further comprising a hardened outer surface, said hardened outer surface produced by a process selected from the group consisting of silver doping and boronation.

47. The medical implant according to claim 2 comprising an expandable stent for supporting a blood, urinary, or gastrointestinal vessel from collapsing inward, the stent comprising:

(a) a radially outwardly expandable substantially cylindrical stent body of a metal alloy comprising:

(i) titanium;
(ii) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (iii) optionally up to about 20 wt% zirconium;

the stent body in its unexpected state sized for ease of insertion into a vessel and in expanded state sized for propping open a vessel, the stent body having a bore therethrough for receiving a means for expanding the body radially outwardly.

48. The stent of claim 47, wherein the metal alloy comprises from about 0.5 to about 20 wt% zirconium.

49. The stent of claim 48, wherein surfaces of the stent are hardened by a process selected from the group consisting of oxygen diffusion hardening, nitrogen hardening, and physical vapour deposition, and chemical vapour deposition.

50. The stent of claim 47, wherein surfaces of the stent are coated with a composition comprising a medicament.

51. The stent of claim 47 further comprising a coating on surfaces of the stent, said coating applied by a process selected from the group consisting of silver doping and boronation.

52. The medical implant according to claim 2 comprising a biocompatible lead or sensor for conducting electrical signals to or from an organ in a living body, the lead comprising:

an elongate flexible body having distal and proximal ends, the flexible body comprising:

(a) an electrically conductive core of a metal alloy comprising:

(i) titanium (ii) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (iii) optionally up to about 20 wt% zirconium;
for carrying the electrical signals.

53. The lead of claim 52, wherein the metal alloy comprises from about 0.5 to about 20 wt% zirconium.

54. The lead of claim 53 further comprising a thin non-electrically conductive layer surrounding the electrically conductive core.

55. The lead of claim 53 further comprising a coating on the non electrically conducting layer, said coating comprising a medicament.

56. The medical implant according to claim 2 comprising a low modulus, biocompatible, percutaneous implant that penetrates the skin of a living body and thereby protrudes from the body, the implant comprising:

(a) a low modulus metallic implant body fabricated from a metal alloy comprising:

(i) titanium;
(ii) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (iii) optionally up to about 20 wt% zirconium;

said implant body having a first end for insertion into said patient and a portion with a second end for extending outside of said patient.

57. The percutaneous implant of claim 56, wherein the metal alloy comprises from about 0.5 to about 20 wt%
zirconium.

58. The percutaneous implant of claim 57 further comprising a hardened outer surface, said hardened outer surface produced by a process selected from the group consisting of oxygen diffusion hardening, nitrogen hardening, physical vapour deposition, and chemical vapour deposition.

59. The percutaneous implant of claim 56, further comprising a coating on surfaces of the implant body, said coating comprising a medicament.

60. The percutaneous implant of claim 56 further comprising a wear-resistant layer on surfaces of the implant body, the wear-resistant layer produced by a process selected from the group consisting of silver doping and boronation.

61. The medical implant according to claim 2 comprising an external mechanical heart including Components with surfaces subject to mechanical wear and microfretting wear, the improvement comprising:

components of a metal alloy comprising:

(a) titanium;
(b) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (c) optionally up to about 20 wt% zirconium.

62. The mechanical heart of claim 61, wherein the metal metal alloy comprises from about 0.5 to about 20 wt% zirconium.

63. The mechanical heart of claim 62 further comprising a hardened outer surface on the components subject to mechanical and microfretting wear.

64. The mechanical heart of claim 61 further comprising a wear-resistant coating on surfaces subject to mechanical wear and microfretting wear, the wear-resistant coating produced by a process selected from the group consisting of silver doping and boronation.

65. The medical implant according to claim 2 comprising an external mechanical heart including components that contact blood when said heart is used to supply blood to a patient, the improvement comprising the components fabricated from a metal alloy comprising:

(a) titanium;
(b) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (c) optionally up to about 20 wt% zirconium.

66. The mechanical heart of claim 65, wherein the metal alloy comprises from about 0.5 to about 20 wt%
zirconium.

67. The mechanical heart of claim 65 further comprising a coating on outer surfaces of the components that contact blood, the coating comprising a medicament.

68. The medical implant according to claim 2 comprising a vascular graft of enhanced durability, crush-resistance, low thrombogenicity, and hemocompatibility, said graft comprising:

(a) titanium;
(b) from about 10 to about 20 wt% niobium or from about 35 to about 50 wt% niobium; and (c) optionally up to about 20 wt% zirconium;

said tubular body having a bore therethrough, said tubular body sized to replace a section of removed blood vessel.

69. The vascular graft of claim 68, wherein the metal alloy comprises from about 0.5 to about 20 wt.%
zirconium.

70. The vascular graft of claim 69 further comprising a hardened surface on the tubular body, said hardened surface produced by a process selected from the group consisting of oxygen diffusion hardening, nitrogen hardening, physical vapour deposition and chemical vapour deposition.

71. The vascular graft of claim 68 further comprising a coating on the tubular body, said coating selected from the group consisting a heparin, phosphatidyl choline, and a medicament.

72. The vascular graft of claim 68, further comprising a wear-resistant coating on the tubular body, said wear-resistant coating applied by a process selected from the group consisting of silver doping and boronation.