U.S. patent application number 12/337055 was filed with the patent office on 2009-04-16 for radiopaque nitinol alloys for medical devices.
This patent application is currently assigned to ABBOTT CARDIOVASCULAR SYSTEMS, Inc.. Invention is credited to John F. Boylan, Daniel L. Cox.
Application Number | 20090098013 12/337055 |
Document ID | / |
Family ID | 25025360 |
Filed Date | 2009-04-16 |
United States Patent
Application |
20090098013 |
Kind Code |
A1 |
Boylan; John F. ; et
al. |
April 16, 2009 |
RADIOPAQUE NITINOL ALLOYS FOR MEDICAL DEVICES
Abstract
A radiopaque nitinol medical device such as a stent for use with
or implantation in a body lumen is disclosed. The stent is made
from a superelastic alloy such as nickel-titanium or nitinol, and
includes a ternary element selected from the group of chemical
elements consisting of iridium, platinum, gold, rhenium, tungsten,
palladium, rhodium, tantalum, silver, ruthenium, or hafnium. The
added ternary element improves the radiopacity of the nitinol stent
comparable to that of a stainless steel stent of the same size and
strut pattern coated with a thin layer of gold. The nitinol stent
has improved radiopacity yet retains its superelastic and shape
memory behavior and further maintains a thin strut/wall thickness
for high flexibility.
Inventors: |
Boylan; John F.; (Murrietta,
CA) ; Cox; Daniel L.; (Palo Alto, CA) |
Correspondence
Address: |
FINNEGAN, HENDERSON, FARABOW, GARRETT & DUNNER;LLP
901 NEW YORK AVENUE, NW
WASHINGTON
DC
20001-4413
US
|
Assignee: |
ABBOTT CARDIOVASCULAR SYSTEMS,
Inc.
|
Family ID: |
25025360 |
Appl. No.: |
12/337055 |
Filed: |
December 17, 2008 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10900632 |
Jul 27, 2004 |
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12337055 |
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09752212 |
Dec 27, 2000 |
6855161 |
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10900632 |
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Current U.S.
Class: |
420/581 ;
420/417; 420/441; 420/580 |
Current CPC
Class: |
Y10T 29/49995 20150115;
A61F 2/915 20130101; A61L 31/022 20130101; A61F 2230/0013 20130101;
A61L 31/18 20130101; A61F 2002/91525 20130101; A61F 2002/91516
20130101; A61L 2400/16 20130101; A61F 2002/91533 20130101; C22C
19/00 20130101; A61F 2002/91575 20130101; Y10S 623/901 20130101;
A61F 2250/0098 20130101 |
Class at
Publication: |
420/581 ;
420/417; 420/441; 420/580 |
International
Class: |
C22C 30/00 20060101
C22C030/00; C22C 14/00 20060101 C22C014/00; C22C 19/03 20060101
C22C019/03 |
Claims
1-18. (canceled)
19. An alloy, comprising: titanium in an amount ranging from about
30 to about 52 atomic percent; platinum in an amount ranging from
greater than 0 to about 15 atomic percent; and the balance
nickel.
20. The alloy of claim 19, wherein platinum is present in an amount
greater than or equal to 2.5 atomic percent.
21. The alloy of claim 20, wherein the platinum is present in an
amount ranging from greater than or equal to 2.5 atomic percent to
less than or equal to about 15 atomic percent.
22. The alloy of claim 19, further comprising a quaternary alloying
element.
23. The alloy of claim 22, wherein said quaternary alloying element
is iron.
24. A radiopaque super-elastic alloy, comprising: about 37.49
atomic percent titanium; about 39.48 atomic percent nickel; and the
balance of platinum.
25. The radiopaque super-elastic alloy of claim 24, further
comprising a quaternary alloying element.
26. The radiopaque super-elastic alloy of claim 25, wherein said
quaternary alloying element is iron.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This is a continuation application of co-pending parent
application having U.S. Ser. No. 09/752,212, filed Dec. 27, 2000,
the entire contents of which are hereby incorporated by
reference.
BACKGROUND OF THE INVENTION
[0002] The present invention generally relates to self-expanding
medical devices. More precisely, the present invention relates to
self-expanding medical devices made of radiopaque nitinol that can
be used in essentially any body lumen. Such devices include
stents.
[0003] Stents are typically implanted in a body lumen, such as
carotid arteries, coronary arteries, peripheral arteries, veins, or
other vessels to maintain the patency of the lumen. These devices
are frequently used in the treatment of atherosclerotic stenosis in
blood vessels especially after percutaneous transluminal
angioplasty (PTA) or percutaneous transluminal coronary angioplasty
(PTCA) procedures with the intent to reduce the likelihood of
restenosis of a vessel. Stents are also used to support a body
lumen, tack-up a flap or dissection in a vessel, or in general
where the lumen is weak to add support.
[0004] During PTCA procedures it is common to use a dilatation
catheter to expand a diseased area to open the patient's lumen so
that blood flows freely. Despite the beneficial aspects of PTCA
procedures and its widespread and accepted use, it has several
drawbacks, including the possible development of restenosis and
perhaps acute thrombosis and sub-acute closure. This recurrent
stenosis has been estimated to occur in seventeen to fifty percent
of patients despite the initial PTCA procedure being successful.
Restenosis is a complex and not fully understood biological
response to injury of a vessel which results in chronic hyperplasia
of the neointima. This neointimal hyperplasia is activated by
growth factors which are released in response to injury. Acute
thrombosis is also a result of vascular injury and requires
systemic antithrombotic drugs and possibly thrombolytics as well.
This therapy can increase bleeding complications at the catheter
insertion site and may result in a longer hospital stay. Sub-acute
closure is a result of thrombosis, elastic recoil, and/or vessel
dissection.
[0005] Several procedures have been developed to combat restenosis
and sub-acute or abrupt closure, one of which is the delivery and
implanting of an intravascular stent. Stents are widely used
throughout the United States and in Europe and other countries.
Generally speaking, the stents can take numerous forms. One of the
most common is a generally cylindrical, hollow tube that holds open
the vascular wall at the area that has been dilated by a dilation
catheter. One highly regarded stent used and sold in the United
States is known under the tradename ACS Multi-Link Stent, which is
made by Advanced Cardiovascular Systems, Inc., Santa Clara,
Calif.
[0006] In expandable stents that are delivered with expandable
catheters, such as balloon catheters, the stents are positioned
over the balloon portion of the catheter and are expanded from a
reduced diameter to an enlarged diameter greater than or equal to
the inner diameter of the arterial wall by inflating the balloon.
Stents of this type can be expanded to an enlarged diameter by
deforming the stent, by engagement of the stent walls with respect
to one another, and by one way engagement of the stent walls
together with endothelial growth onto and over the stent.
[0007] Examples of intravascular stents can be found in U.S. Pat.
No. 5,292,331 (Boneau); U.S. Pat. No. 4,580,568 (Gianturco); U.S.
Pat. No. 4,856,516 (Hillstead); U.S. Pat. No. 5,092,877 (Pinchuk);
and U.S. Pat. No. 5,514,154 (Lau et al.), which are incorporated
herein by reference in their entirety.
[0008] The problem with some prior art stents, especially those of
the balloon expandable type, is that they are often stiff and
inflexible. These balloon expandable type stents are commonly
formed from stainless steel alloys and the stents are constructed
so that they are expanded beyond their elastic limit. As a result,
such stents are permanently deformed by the inflation balloon
beyond their elastic limits to hold open a body lumen and thus
maintain patency of that body lumen. There are several commercially
available balloon expandable stents that are widely used; they are
generally implanted in the coronary arteries after a PTCA procedure
mentioned earlier.
[0009] Stents are often times implanted in vessels that are closer
to the surface of the body, such as in the carotid arteries in the
neck or in peripheral arteries and veins in the leg. Because these
stents are so close to the surface of the body, they are
particularly vulnerable to impact forces that can partially or
completely collapse the stent and thereby block fluid flow in the
vessel. Other forces can impact balloon expandable stents and cause
similar partial or total vessel blockage. For instance, under
certain conditions, muscle contractions might also cause balloon
expandable stents to collapse partially or completely. The collapse
occludes the lumen and restricts blood flow in the vessel in which
they are implanted.
[0010] Since balloon expandable stents are plastically deformed,
once collapsed or crushed they remain so, permanently blocking the
vessel. Thus, balloon expandable stents under certain conditions
might pose an undesirable condition for the patient.
[0011] Self-expanding stents as the name implies self-expand
through the properties of the material constituting the stent. The
inflation force of a balloon catheter is usually not necessary to
deploy this kind of stent.
[0012] Important applications including those mentioned above have
prompted designers to seek out superelastic shape memory alloys to
exploit the materials' properties in their self-expanding stents.
Examples of applying superelastic nickel-titanium alloys to a
self-expanding stent and other medical devices are disclosed in
U.S. Pat. Nos. 4,665,906; 5,067,957; 5,190,546; and 5,597,378 to
Jervis and U.S. Pat. No. 4,503,569 to Dotter. Another example is
disclosed in European Patent Application Publication No.
EP0873734A2, entitled "Shape Memory Alloy Stent." This publication
suggests a stent for use in a lumen in a human or animal body
having a generally tubular body formed from a shape memory alloy
which has been treated so that it exhibits enhanced elastic
properties. The publication further suggests use of specified
ternary elements in a nickel-titanium alloy to obtain desired
engineering characteristics.
[0013] Use of a ternary element in a superelastic stent is also
shown in, for example, U.S. Pat. No. 5,907,893 to Zadno-Azizi et
al. As a general proposition, there have been attempts at adding a
ternary element to nickel-titanium alloys as disclosed in, for
instance, U.S. Pat. No. 5,885,381 to Mitose et al
[0014] Clearly, self-expanding, nickel-titanium stents are useful
and valuable to the medical field. But a distinct disadvantage with
self-expanding nickel-titanium stents is the fact that they are not
sufficiently radiopaque as compared to a comparable structure made
from gold or tantalum. For example, radiopacity permits the
cardiologist or physician to visualize the procedure involving the
stent through use of fluoroscopes or similar radiological
equipment. Good radiopacity is therefore a useful feature for
self-expanding nickel-titanium stents to have.
[0015] Radiopacity can be improved by increasing the strut
thickness of the nickel-titanium stent. But increasing strut
thickness detrimentally affects the flexibility of the stent, which
is a quality necessary for ease of delivery. Another complication
is that radiopacity and radial force co-vary with strut thickness.
Also, nickel-titanium is difficult to machine and thick struts
exacerbates the problem.
[0016] Radiopacity can be improved through coating processes such
as sputtering, plating, or co-drawing gold or similar heavy metals
onto the stent. These processes, however, create complications such
as material compatibility, galvanic corrosion, high manufacturing
cost, coating adhesion or delamination, biocompatibility, loss of
coating integrity following collapse and deployment of the stent,
etc.
[0017] Radiopacity can also be improved by alloy addition. One
specific approach is to alloy the nickel-titanium with a ternary
element. What has been needed and heretofore unavailable in the
prior art is a superelastic nickel-titanium stent that includes a
ternary element to increase radiopacity yet preserves the
superelastic qualities of the nitinol.
SUMMARY OF THE INVENTION
[0018] The present invention relates to a radiopaque medical
device, such as a stent, for use or implantation in a body lumen.
In a preferred embodiment, a radiopaque medical device, such as a
stent, is constructed from a tubular-shaped body having a thin wall
defining a strut pattern; wherein the tubular body includes a
superelastic, nickel-titanium alloy, and the alloy further includes
a ternary element selected from the group of chemical elements
consisting of iridium, platinum, gold, rhenium, tungsten,
palladium, rhodium, tantalum, silver, ruthenium, or hafnium. In a
preferred embodiment, the stent according to the present invention
has 42.8 atomic percent nickel, 49.7 atomic percent titanium, and
7.5 atomic percent platinum.
[0019] As a result, the present invention stent is highly
radiopaque as compared to an identical structure made of medical
grade stainless steel that is coated with a thin layer of gold.
From another perspective, for a given stent having a certain level
of radiopacity, the present invention stent having identical
dimensions and strut pattern has at least a 10 percent reduction in
strut thickness yet maintains that same level of radiopacity.
[0020] Self-expanding nitinol stents are collapsed (that is,
loaded) and then constrained within a delivery system. At the point
of delivery, the stent is released (that is, unloaded) and allowed
to return to its original diameter. The stent is designed to
perform various mechanical functions within the lumen, all of which
are based upon the lower unloading plateau stress. Therefore, it is
crucial that the ternary element alloyed with the binary
nickel-titanium does not diminish the superelastic characteristics
of the nickel-titanium.
[0021] To achieve the sufficient degree of radiopacity yet
maintaining the superelastic engineering properties of a binary
nickel-titanium, preferably, the radiopaque stent of the present
invention includes platinum whose atomic percent is greater than or
equal to 2.5 and less than or equal to 15. In an alternative
embodiment, the nickel-titanium is alloyed with palladium whose
atomic percent is greater than or equal to 2.5 and less than or
equal to 20. With such compositions, the stress-strain hysteresis
curve of the present invention radiopaque nitinol alloy closely
approximates the idealized stress-strain hysteresis curve of binary
nickel-titanium.
[0022] The present invention further contemplates a method for
providing a radiopaque nitinol stent. In a preferred embodiment,
the method entails providing a tubular-shaped body having a thin
wall, wherein the body includes a superelastic nickel-titanium
alloy and the alloy further includes a ternary element selected
from the group of chemical elements consisting of iridium,
platinum, gold, rhenium, tungsten, palladium, rhodium, tantalum,
silver, ruthenium, or hafnium; forming a strut pattern wherein the
stent is highly radiopaque. The step of providing a tubular-shaped
body includes melting nickel, titanium, and the ternary element and
cooling the mixture to form an alloy ingot, hot forming the alloy
ingot, hot or cold forming the alloy ingot into a cylinder,
drilling the cylinder to form tubing, cold drawing the tubing, and
annealing the tubing.
[0023] The present invention of course envisions the minor addition
of a quaternary element, for example, iron, to further enhance the
alloy's formability or its thermomechanical properties. In short,
the presence of elements in addition to the ternary elements cited
above is contemplated.
[0024] In a preferred embodiment, an austenite finish temperature
(A.sub.f) of the superelastic alloy in the stent is greater than or
equal to zero and less than or equal to 37 degrees C. Also in the
preferred embodiment, the ingot after melting includes an austenite
finish temperature (A.sub.f) of greater than or equal to 0 degrees
C. and less than or equal to 40 degrees C. The tubing includes an
austenite finish temperature (A.sub.f) of greater than or equal to
-15 degrees C. and less than or equal to 15 degrees C.
[0025] Other features and advantages of the present invention will
become more apparent from the following detailed description of the
invention when taken in conjunction with the accompanying exemplary
drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
[0026] FIG. 1 is a side elevational view, partially in section,
depicting a stent mounted on a delivery catheter and expanded
within a damaged vessel, pressing a damaged vessel lining against
the vessel wall.
[0027] FIG. 2 is a side elevational view, partially in section,
depicting an expanded stent within the vessel after withdrawal of
the delivery catheter.
[0028] FIG. 3 is an idealized stress-strain hysteresis curve for a
superelastic material.
[0029] FIG. 4 is a plan view of the flattened strut pattern of an
exemplary embodiment superelastic stent.
[0030] FIG. 5 is a group of empirical data curves illustrating the
highly similar stress-strain relationships among binary nitinol and
the nickel-titanium-palladium and nickel-titanium-platinum alloys
used in the present invention.
DETAILED DESCRIPTION OF THE INVENTION
[0031] The present invention relates to a medical device made of
radiopaque nitinol. For the sake of illustration, the following
exemplary embodiments are directed to stents, although it is
understood that the present invention is applicable to other
medical devices usable in a body lumen as well.
[0032] The stents of the present invention can have virtually any
configuration that is compatible with the body lumen in which they
are implanted. The stent should preferably be configured so that
there is a substantial amount of open area and preferably the open
area to metal ratio is at least 80 percent. The stent should also
be configured so that dissections or flaps in the body lumen wall
are covered and tacked up by the stent.
[0033] Referring to FIGS. 1, 2, and 4, in a preferred embodiment, a
stent 10 of the present invention is formed partially or completely
of alloys such as nitinol (NiTi) which have superelastic (SE)
characteristics. Stent 10 is somewhat similar to the stent
disclosed in U.S. Pat. No. 5,569,295, "Expandable Stents and Method
for Making Same," issued to Lam on Oct. 29, 1996, which patent is
incorporated herein by reference. Some differences of the present
invention stent from that disclosed in the '295 patent is that the
present invention stent is preferably constructed of a superelastic
material with the addition of a ternary element, and the strut
pattern has changed. Of course, the configuration of the stent 10
is just one example of many stent configurations that are
contemplated by the present invention.
[0034] Turning to FIG. 4, stent 10 has a tubular form which
preferably includes a plurality of radially expandable cylindrical
elements 24 disposed generally coaxially and interconnected by
members 26 disposed between adjacent cylindrical elements 24. The
shapes of the struts 12 forming the strut pattern are designed so
they can preferably be nested. This strut pattern is best seen from
the flattened plan view of FIG. 4. The serpentine patterned struts
12 are nested such that the extended portions of the struts of one
cylindrical element 24 intrude into a complementary space within
the circumference of an adjacent cylindrical element. In this
manner, the plurality of cylindrical elements 24 can be more
tightly packed lengthwise.
[0035] As introduced above, an exemplary stent of the present
invention includes a superelastic material. In a general sense,
superelasticity implies that the material can undergo a large
degree of reversible strain as compared to common steel. In a
technical sense, the term "superelasticity" and sometimes
"pseudoelasticity" refer to an isothermal transformation in
nitinol. More specifically, it refers to stress inducing a
martensitic phase from an austenitic phase. Alloys having
superelastic properties generally have at least two phases: a
martensitic phase, which has a relatively low tensile strength and
which is stable at relatively low temperatures, and an austenitic
phase, which has a relatively high tensile strength and which is
stable at temperatures higher than the martensitic phase.
Superelastic characteristics generally allow the metal stent to be
deformed by collapsing the stent and creating stress which causes
the NiTi to reversibly change to the martensitic phase. The stent
is restrained in the deformed condition inside a delivery sheath
typically to facilitate the insertion into a patient's body, with
such deformation causing the isothermal phase transformation. Once
within the body lumen, the restraint on the stent is removed,
thereby reducing the stress thereon so that the superelastic stent
returns towards its original undeformed shape through isothermal
transformation back to the austenitic phase. Under these
conditions, the stent can be described as self-expanding.
[0036] Returning to FIG. 1, the graphic illustrates, in a partial
cross-sectional view, the distal end of a rapid exchange stent
delivery system that includes a guide wire 14, a delivery sheath
16, and an intravascular catheter 18. For the sake of clarity, the
illustration of the delivery system in FIG. 1 has been simplified.
It is just one example of a delivery system that may be used with
the present invention. More details of a delivery system
specifically for use with a self-expanding stent may be found in,
for example, U.S. Pat. No. 6,077,295 to Limon et al., entitled
"Self-Expanding Stent Delivery System," which is incorporated
herein by reference. Other delivery systems such as over-the-wire
may be used without departing from the scope of the instant
invention.
[0037] FIG. 1 further shows an optional expandable balloon 20
inflated through an inflation lumen (not shown), although the
balloon is typically not needed for a self-expanding stent. The
stent 10 is first crimped on to the deflated balloon 20, and the
entire assembly is kept underneath the delivery sheath 16 until the
moment the stent 10 is deployed. The stent 10 is self-expanding so
that when the sheath 16 is withdrawn, the stent 10 expands to its
larger deployment diameter without assistance from the balloon 20.
Nevertheless, some procedures specifically use the balloon 20 to
further expand the stent 10 for improved seating in the artery wall
29.
[0038] FIG. 2 illustrates the self-expanding stent 10 in the
expanded condition after the delivery system has been removed. If
an external force is applied to the artery 28, the expanded stent
10 temporarily and at least partially collapses or deforms. As the
stent 10 deforms, stress in the nickel-titanium alloy causes an
isothermal phase transformation from the austenitic phase to the
martensitic phase. When the external force is removed, the stress
in stent 10 is likewise diminished so that the stent quickly
transforms back from the martensitic phase to the austenitic phase.
As this almost instantaneous, isothermal transformation occurs, the
stent 10 returns to its fully expanded state and the artery remains
open. When the superelastic stent 10 is implanted in an artery 28,
its high resilience effectively maintains the patency of the artery
while minimizing the risk of permanent arterial collapse at the
implant site if the stent is temporarily deformed due to external
forces. Furthermore, the resilience of the stent 10 supports the
flap 30 to maintain patency of the artery.
[0039] Stent 10 is preferably formed from a superelastic material
such as nickel-titanium and undergoes an isothermal transformation
when stressed if in the austenitic phase. For most purposes, the
transformation temperature for the stent 10 is preferably set low
enough such that the nickel-titanium alloy is in the austenitic
phase while at body temperature.
[0040] According to theory, when stress is applied to a specimen of
a metal such as nitinol exhibiting superelastic characteristics at
a temperature at or above that which the transformation of the
martensitic phase to the austenitic phase is complete, the specimen
deforms elastically until it reaches a particular stress level
where the alloy then undergoes a stress-induced phase
transformation from the austenitic phase to the martensitic phase.
As the phase transformation progresses, the alloy undergoes
significant increases in strain with little or no corresponding
increases in stress. The strain increases while the stress remains
essentially constant until the transformation of the austenitic
phase to the martensitic phase is complete. Thereafter, further
increase in stress is necessary to cause further deformation. The
martensitic metal first yields elastically upon the application of
additional stress and then plastically with permanent residual
deformation.
[0041] If the load on the specimen is removed before any permanent
deformation has occurred, the stress-induced martensite elastically
recovers and transforms back to the austenitic phase. The reduction
in stress first causes a decrease in strain. As stress reduction
reaches the level at which the martensitic phase begins to
transform back into the austenitic phase, the stress level in the
specimen remains essentially constant (but less than the constant
stress level at which the austenitic crystalline structure
transforms to the martensitic crystalline structure until the
transformation back to the austenitic phase is complete); i.e.,
there is significant recovery in strain with only negligible
corresponding stress reduction. After the transformation back to
austenite is complete, further stress reduction results in elastic
strain reduction. This ability to incur significant strain at
relatively constant stress upon the application of a load and to
recover from the deformation upon the removal of the load is
commonly referred to as "superelasticity" and sometimes
"pseudoelasticity."
[0042] FIG. 3 illustrates an idealized stress-strain hysteresis
curve for a superelastic, binary nickel-titanium alloy. The
relationship is plotted on x-y axes, with the x axis representing
strain and the y axis representing stress. For ease of
illustration, the x-y axes are labeled on a scale typical for
superelastic nitinol, with stress from 0 to 60 ksi and strain from
0 to 9 percent, respectively.
[0043] Looking at the plot in FIG. 3, the line from point A to
point B represents the elastic deformation of the nickel-titanium
alloy. After point B the strain or deformation is no longer
proportional to the applied stress and it is in the region between
point B and point C that the stress-induced transformation of the
austenitic phase to the martensitic phase begins to occur.
[0044] At point C moving toward point D, the material enters a
region of relatively constant stress with significant deformation
or strain. This constant or plateau region is known as the loading
stress, since it represents the behavior of the material as it
encounters continuous increasing strain. It is in this plateau
region C-D that the transformation from austenite to martensite
occurs.
[0045] At point D the transformation to the martensitic phase due
to the application of stress to the specimen is substantially
complete. Beyond point D the martensitic phase begins to deform,
elastically at first, but, beyond point E, the deformation is
plastic or permanent.
[0046] When the stress applied to the superelastic metal is
removed, the material behavior follows the curve from point E to
point F. Within the E to F region, the martensite recovers its
original shape, provided that there was no permanent deformation to
the martensitic structure. At point F in the recovery process, the
metal begins to transform from the stress-induced, unstable,
martensitic phase back to the more stable austenitic phase.
[0047] In the region from point G to point H, which is also an
essentially constant or plateau stress region, the phase
transformation from martensite back to austenite takes place. This
constant or plateau region G-H is known as the unloading stress.
The line from point I to the starting point A represents the
elastic recovery of the metal to its original shape.
[0048] Binary nickel-titanium alloys that exhibit superelasticity
have an unusual stress-strain relationship as just described and as
plotted in the curve of FIG. 3. As emphasized above, the
superelastic curve is characterized by regions of nearly constant
stress upon loading, identified above as loading plateau stress C-D
and unloading plateau stress G-H. Naturally, the loading plateau
stress C-D always has a greater magnitude than the unloading
plateau stress G-H. The loading plateau stress represents the
period during which martensite is being stress-induced in favor of
the original austenitic crystalline structure. As the load is
removed, the stress-induced martensite transforms back into
austenite along the unloading plateau stress part of the curve. The
difference in stress between the stress at loading C-D and
unloading stress G-H defines the hysteresis of the system.
[0049] The present invention seeks to preserve the superelastic
qualities of nickel-titanium alloys just described yet improve upon
the material's radiopacity by addition of a ternary element. This
is preferably accomplished in one embodiment by forming a
composition consisting essentially of about 30 to about 52 percent
titanium and the balance nickel and up to 10 percent of one or more
additional ternary alloying elements. Such ternary alloying
elements may be selected from the group consisting of iridium,
platinum, gold, rhenium, tungsten, palladium, rhodium, tantalum,
silver, ruthenium, or hafnium. In the preferred embodiment, the
atomic percentage of platinum is greater than or equal to 2.5 and
less than or equal to 15. In an alternative embodiment, the atomic
percentage of palladium is greater than or equal to 2.5 and less
than or equal to 20.
[0050] A preferred embodiment stent according to the present
invention has 42.8 atomic percent nickel, 49.7 atomic percent
titanium, and 7.5 atomic percent platinum. Through empirical
studies, the aforementioned compositions produce stent patterns
having a radiopacity comparable to the same size and pattern stent
made from 316 L stainless steel with a 2.7 to 6.5 .mu.m gold
coating.
[0051] In various alternative embodiments, the present invention
contemplates the minor addition of a quaternary element, for
example, iron, to further enhance the alloy's formability or its
thermomechanical properties. The presence of impurities such as
carbon or oxygen or the like in the present invention alloy is also
possible.
[0052] A preferred method of fabricating the present invention
superelastic, radiopaque metallic stent entails first fashioning
nickel-titanium tubing. The tubing is made from vacuum induction
melting nickel and titanium with the ternary element according to
the compositions suggested above. The ingot is then remelted for
consistency. The ingot is next hot rolled into bar stock, then
straightened and sized, and hot or cold formed into a cylinder. The
cylinder is gun drilled to form the tubing. Instead of gun
drilling, other methods of material removal known in the art may be
used, including electric discharge machining (EDM), laser beam
machining, and the like. Next, the tubing is cold drawn and
annealed repeatedly to achieve the finished dimensions.
[0053] Any of the foregoing preferred embodiment steps may be
repeated, taken out of sequence, or omitted as necessary depending
on desired results. From here on, the tubing follows conventional
stent fabrication techniques such as laser cutting the strut
pattern, heat setting, etc.
[0054] The following are additional processing guide posts for the
present invention to achieve a sufficiently radiopaque stent yet
maintaining the superelastic stress-strain behavior of the alloy.
Empirical evidence suggests that, in various preferred embodiments,
a Ni--Ti--Pd or Ni--Ti--Pt ingot should have the following
austenite finish temperature: 0 degrees C..ltoreq.A.sub.f.ltoreq.40
degrees C. The Ni--Ti--Pd or Ni--Ti--Pt tubing should exhibit an
austenite finish temperature of: -15 degrees
C..ltoreq.A.sub.f.ltoreq.15 degrees C. In an exemplary embodiment,
the final laser cut Ni--Ti--Pd or Ni--Ti--Pt stent should exhibit
an austenite finish temperature of: 0 degrees
C..ltoreq.A.sub.f.ltoreq.37 degrees C. Of course, the A.sub.f of
the finished laser cut stent can be set as needed by various heat
treating processes known in the art.
[0055] It is understood that the austenite finish temperature
(A.sub.f) is defined to mean the temperature at which the material
completely reverts to austenite. In technical terms, the A.sub.f
(and other transformation temperatures A.sub.S, M.sub.S, M.sub.f)
as it applies to an ingot made of Ni--Ti--Pd or Ni--Ti--Pt, for
example, is determined by a Differential Scanning Calorimeter (DSC)
test, known in the art. The DSC test method to determine
transformation temperatures for the ingot is guided by ASTM
standard no. F 2004-00, titled "Standard Test Method For
Transformation Temperature Of Nickel-Titanium Alloys By Thermal
Analysis."
[0056] The "active A.sub.f" for the tubing and the finished stent
is determined by a bend and free recovery test, also known in the
art. In such a test, the tubing is cooled to under the M.sub.f
temperature, deformed, and warmed up. While monitoring the
increasing temperature, the point of final recovery of the
deformation in the tubing approximates the A.sub.f of the material.
The active A.sub.f testing technique is guided by a second ASTM
standard entitled "Standard Test Method For Determination Of
Transformation Temperature Of Nickel-Titanium Shape Memory Alloys
By Bend And Free Recovery," or by equivalent test methods known in
the art.
[0057] Samples of wire made in accordance with the foregoing
exemplary embodiments were tested. Specifically, the stress-strain
relationship based on empirical data for nickel-titanium-palladium
and nickel-titanium-platinum are plotted against binary nitinol in
FIG. 5. Curve A corresponds to a sample of
nickel-titanium-platinum. Curve B is based on a sample of binary
nitinol. Curve C is based on a sample of nickel-titanium-palladium.
To generate the empirical data, the wire samples were placed under
increasing tension until past the phase transformation from their
initial austenitic phase to their martensitic phase. Tension was
then slowly released prior to any plastic deformation until stress
on the samples dropped to zero with full deformation recovery.
[0058] As is apparent from the plot of FIG. 5, the present
invention nickel-titanium-palladium and nickel-titanium-platinum
alloys have stress-strain curves that closely follow the hysteresis
curve for binary nitinol. All three curves have essentially flat
loading and unloading plateau stresses indicating the presence of a
phase transformation that is characteristic of superelastic metals.
Hence, the present invention nitinol stent incorporates a ternary
element, in these exemplary embodiments palladium or platinum, to
improve radiopacity yet the materials' superelastic capability is
preserved. What has been missing heretofor is empirical evidence
that this level of radiopacity can be achieved while preserving the
superelastic characteristics of these alloys.
[0059] The present invention further provides a nitinol stent
having improved radiopacity without reliance on increasing the
stent wall thickness or strut thickness. Increasing wall or strut
thicknesses detracts from the flexibility of the stent, which is
detrimental to deliverability. Rather, the present invention
superelastic nitinol stent has a thin wall/strut thickness and/or
strut cross-sectional area akin to a conventional stainless steel
stent, and has comparable radiopacity to a stainless steel stent
with a thin coating of gold. The wall/strut thickness is defined by
the difference between the inside diameter and the outside diameter
of the tube.
[0060] Indeed, the improved radiopacity of the present invention
stent can be characterized strictly by strut thickness. In this
con, the present invention radiopaque stent has a reduced strut
thickness yet exhibits the radiopacity of an identical stent having
thicker struts. In other words, given a stent exhibiting a certain
level of radiopacity, the present invention stent having the
identical dimensions and strut pattern achieves that level of
radiopacity yet it has at least a 10 percent reduction in strut
thickness as compared to the reference stent.
[0061] Alternatively, the 10 percent reduction can also be
quantified in terms of the cross-sectional area of the strut. That
is, for a given stent having a certain level of radiopacity with
struts with a given cross-sectional area, the present invention
stent having the same dimensions and strut pattern achieves the
same level of radiopacity but has struts with at least a 10 percent
reduction in cross-sectional area as compared to the reference
stent.
[0062] Another aspect of nitinol aside from its superelasticity is
shape memory. The present invention can also be employed with
respect to this physical attribute as described below.
[0063] The shape memory effect allows a nitinol structure to be
deformed to facilitate its insertion into a body lumen or cavity,
and then heated within the body so that the structure returns to
its original, set shape. Nitinol alloys having shape memory effect
generally have at least two phases: a martensitic phase, which has
a relatively low tensile strength and which is stable at relatively
low temperatures, and an austenitic phase, which has a relatively
high tensile strength and which is stable at temperatures higher
than the martensitic phase.
[0064] Shape memory effect is imparted to the alloy by heating the
nickel-titanium metal to a temperature above which the
transformation from the martensitic phase to the austenitic phase
is complete; i.e., a temperature above which the austenitic phase
is stable. The shape of the metal during this heat treatment is the
shape "remembered." The heat-treated metal is cooled to a
temperature at which the martensitic phase is stable, causing the
austenitic phase to transform to the martensitic phase. The metal
in the martensitic phase is then plastically deformed, e.g., to
facilitate the entry thereof into a patient's body. Subsequent
heating of the deformed martensitic phase to a temperature above
the martensite to austenite transformation temperature causes the
deformed martensitic phase to transform to the austenitic phase.
During this phase transformation the metal reverts back towards its
original shape.
[0065] The recovery or transition temperature may be altered by
making minor variations in the composition of the metal and in
processing the material. In developing the correct composition,
biological temperature compatibility must be determined in order to
select the correct transition temperature. In other words, when the
stent is heated, it must not be so hot that it is incompatible with
the surrounding body tissue. Other shape memory materials may also
be utilized, such as, but not limited to, irradiated memory
polymers such as autocrosslinkable high density polyethylene
(HDPEX). Shape memory alloys are known in the art and are discussed
in, for example, "Shape Memory Alloys," Scientific American, Vol.
281, pp. 74-82 (November 1979), incorporated herein by
reference.
[0066] Shape memory alloys undergo a transition between an
austenitic phase and a martensitic phase at certain temperatures.
When they are deformed while in the martensitic phase, they retain
this deformation as long as they remain in the same phase, but
revert to their original configuration when they are heated to a
transition temperature, at which time they transform to their
austenitic phase. The temperatures at which these transitions occur
are affected by the nature of the alloy and the condition of the
material. Nickel-titanium-based alloys (NiTi), wherein the
transition temperature is slightly lower than body temperature, are
preferred for the present invention. It is desirable to have the
transition temperature set at just below body temperature to insure
a rapid transition from the martensitic state to the austenitic
state when the stent is implanted in a body lumen.
[0067] Turning again to FIGS. 1, 2, and 4, the present invention
stent 10 is formed from a shape memory alloy, such as NiTi
discussed above. After the stent 10 is inserted into an artery 28
or other vessel, the delivery sheath 16 is withdrawn exposing the
stent 10 to the ambient environment. The stent 10 then immediately
expands due to contact with the higher temperature within artery 28
as described for devices made from shape memory alloys. An optional
expandable balloon 20 may be inflated by conventional means to
further expand the stent 10 radially outward.
[0068] Again, if an external force is exerted on the artery, the
stent 10 temporarily at least partially collapses. But the stent 10
then quickly regains its former expanded shape due to its shape
memory qualities. Thus, a crush-resistant stent, having shape
memory characteristics, is implanted in a vessel. It maintains the
patency of a vessel while minimizing both the risk of permanent
vessel collapse and the risk of dislodgment of the stent from the
implant site if the stent is temporarily deformed due to external
forces.
[0069] When the stent 10 is made in accordance with the present
invention, it is also highly radiopaque. The same alloying
processes described earlier are used here to add the ternary
element to increase the radiopacity of the stent. Insofar as the
martensitic to austenitic phase transformation is thermally driven,
the deployment of the present invention stent can be explained in
terms of the shape memory effect.
[0070] While the present invention has been illustrated and
described herein in terms of a radiopaque nitinol stent, it is
apparent to those skilled in the art that the present invention can
be used in other instances. Other modifications and improvements
may be made without departing from the scope of the present
invention.
* * * * *