Note: Descriptions are shown in the official language in which they were submitted.
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METHOD FOR PRODUCING STRAIN INDUCED AUSTENITE
TECHNICAL FIELD
[00011 The present disclosure relates to the processing of nickel-titanium (Ni-
Ti)
ternary alloys and, in particular, to the conditioning of these alloys in a
manner that
enhances the use of such alloys in medical applications.
BACKGROUND
[00031 Metallic engineering materials such as stainless steels, cobalt
chromium alloys
(Co-Cr alloys) and nickel-titanium alloys (Ni-Ti alloys) are used in variety
of medical
device applications. One example of a Ni-Ti alloy is NITINOL. The use of these
alloys
combine mechanical, fatigue, corrosion resistance and biocompatibility
properties to create
devices useful in a number of medical procedures.
[00041 Since the advent of minimally invasive procedures, engineering
designers have
been trying to work with specific geometries whereby metallic components can
be inserted
through very small openings in the human body, routed to the desired location
and then
deployed to a useful size to fulfill the needs of the end application. One
such well-known
device is a coronary stent, a tubular structure used to hold open blocked or
collapsed
arteries. The usual method of getting stents into the space is to collapse the
metal
structure onto a delivery catheter having a sufficiently small overall
diameter so it can be
routed percutaneously to the coronary artery and expanded to a much larger
diameter than
the original insertion diameter.
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[0005] Conventional alloys used in various medical instruments have relied on
stainless steel, complex cobalt chrome alloys (such as Elgiloy TM or L605) all
of which can
have their mechanical properties (i.e. yield strength, ultimate tensile
strength break
strength, etc.) modified through work hardening and annealing. These metals,
even with
very high yield strength, cannot sustain strains much greater than 0.2 percent
(%) without
suffering a permanent set. Once a bend or kink has been sustained in a medical
instrument
or device fabricated from one of the above alloys it is virtually impossible
to straighten and
remove from the body. For many permanent implants (such as stents), the device
may not
need to be removed and the permanent deformation may actually be useful to
keep the
structure in place. However, in the foregoing alloys, Hook's law dictates that
the force to
deploy the implant will increase at a linear rate until the material yield
point is reached and
then the force will continue to increase until the material break point is
reached.
Additionally, these materials have significant spring back after receiving a
significant
deformation.
[0006] Recently medical device engineers have begun designing metallic
components
with shape memory alloys. In general, shape memory alloys such as NiTINoL
having the
proper transformation temperature and processing could potentially offer two
modes of
shape recovery for metallic components inserted into the human body: (1)
superelasticity
and (2) shape memory recovery.
[0007] In the case of superelastic NITINOL the complete "elastic" recovery of
strains
up to 10% due to stress induced martensite (SIM) can be achieved. When
superelastic
NITItvor. components are subjected to a stress, the strain is accommodated by
austenite to
martensite crystalline transformation, rather than by the mechanisms that
prevail in other
alloys such as slip, grain boundary sliding and dislocation motion.
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[0008] Under typical process condition the stress required to form martensite
will be >
60,000 pounds per square inch (psi) (414 mega pascals (MPa) and the reverse
transformation stress will be >30,000 psi (207 MPa). It can be observed that
the reversion
stress is lower than the stress at which martensite forms. These stresses are
referred to as
the upper and lower plateau stresses and their magnitude is dependent on the
alloy
composition, cold working and thermal treatment that the NiTnvOL has received.
As the
temperature of the specimen is raised, the stress magnitude required to
produce SIM is
increased; however when the specimen reaches a critical temperature above (the
Austenite
finish temperature) Af, designated as Md, stress induced martensite cannot be
formed, no
matter how high the stress. In practical applications, this behavior gives
rise to a
limitation on using the super-elastic property since it limits the temperature
range over
which super-elasticity is observed; typically in the binary Ni-Ti alloys, this
is a temperature
range of about 60 Celcius (C)(108 Fahrenheit) (F), although a 40 C (72 F)
range is
more typical. The desirable temperature range for medical and orthodontic
applications is
in the region of body temperature, +10 C to +40 C, can be achieved in these
alloys.
[0009] Others have applied superelastic NiTtrrOL to medical devices using a
50.8
atomic percent (at.%) nickel/balance titanium formulation which has been cold
worked
followed by a low temperature anneal to give a combination of shape memory
and/or
superelastic characteristics. For starting materials having an ingot Af of 0
C, this
processing gives a component with an elastic range of approximately 2% to 8%
over a
temperature range of+15 C to +40 C. However, the shortcomings for deployment
of a
stent application include: (a) the unnecessary bulk of the stent delivery
system (since the
delivery system must resist the high outward radial force of the compressed
stent during
shipping, storage and deployment); (b) the high outward radial force of the
compressed
stent pressing on the inside surface of the delivery sheath can add unwanted
friction during
deployment of the stent from the sheath; (c) at deployment the rapid stent
expansion to its
memorized shape can traumatize the vessel wall; and (c) the stent can cause a
chronic
outward force once deployed that can cause further trauma.
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[0010] Jervis, U.S. Patent No. 5,067,957, discloses that a medical device
component
made from superelastic NITINOL can be externally constrained outside the body
via the
stress induced martensite mechanism, then placed in the body and de-
constrained for
deployment.
[0011] Duerig, et al., U.S. Patent No. 6,312,455, discloses a superelastic
NiUvOL
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 with a point of inflection in the stress-strain curve on
loading. This
enables the body to be deformed inwardly to a transversely compressed
configuration for
insertion into the lumen and then revert towards its initial configuration,
into contact with
and to support the lumen. The shape memory alloy comprises nickel, titanium
and from
about 3 at.% to about 20 at.%, of the alloy composition, of a ternary element
selected
from the group consisting of niobium, hafnium, tantalum, tungsten and gold.
The ratio of
the stress on loading to the stress on unloading at the respective inflection
points on the
loading and unloading curves is at least about 2.5:1, and the difference
between the
stresses on loading and unloading at the inflection points is at least about
250 MPa.
[0012] Besselink, et al., U.S. Patent No. 6,428,634, discloses a method of
processing
a highly elastic stent made from a Ni-Ti-Nb based alloy which contains from
about 4 to
about 14 at.% Nb and in which the atomic percent ratio Ni to Ti is from about
3.8 to 1.2,
comprising working the alloy sufficiently to impart a textured structure to
the alloy, at a
temperature below the recrystallization temperature of the alloy. Preferably,
the alloy is
worked at least 10%, by a technique such as rolling or drawing, or another
technique
which produces a similar crystal structure. The alloy has increased stiffness
compared
with Ni-Ti binary alloys with superelastic properties.
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[0013] For the case of shape memory recovery mentioned earlier, the
thermoelastic
shape memory alloys can change from martensite to austenite and back again on
heating
and cooling over a very small temperature range, typically from 18 C to 55 C.
On
cooling from the austenitic phase, often called the parent phase, martensite
starts to form
at a temperature designated as M, (martensite start) and upon reaching the
lower
temperature, Mf (martensite finish), the alloy is completely martensitic. Upon
heating
from below the Mf temperature the martensite starts to revert to the
austenitic structure at
A,, and when the temperature designated as Af is reached, the alloy is
completely
austenitic. These two crystalline phases have very different mechanical
properties: the
Young's Modulus of austenite is 12 x 106 psi (82,728 MPa), while that for
martensite is
about 4 x 106 psi (27576 MPa); and the yield strength, which depends on the
amount of
cold work the alloy is given, ranges from 28 to 100 thousand pound per square
inch (ksi)
(193 to 689 MPa) for austenite and from 10 to 20 ksi (68 to 138 MPa) for
martensite.
100141 Additionally, a NITINoL structure processed to exhibit shape memory and
deformed in the martensitic state can recover up to 8% strain on heating to
austenite. This
would be an extremely handy way to deploy devices or recover accidental
bending and
kinking of devices in the human body if it were not for the heating and
cooling extremes
that must be achieved.
[0015] Simpson, et al., U.S. Patent No. 4,770,725, discloses a Ni-Ti-Nb shape
memory alloy and article, wherein niobium varies from about 2.5 to 30 at.%.
Also
disclosed is an article made from these nickel/ titanium/niobium alloys.
[00161 Simpson, et al., U.S. Patent No. 4,631,094, discloses a method of
processing a
nickel/titanium-based shape memory alloy. The method comprises over deforming
the
alloy so as to cause at least some amount of non-recoverable strain,
temporarily expanding
the transformation hysteresis by raising the austenite transformation
temperature,
removing the applied stress and then storing the alloy at a temperature less
than the new
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austenite transition temperature. Simpson also discloses an article produced
from this
method.
[0017] Wu, et al., U.S. Patent No. 6,053,992, discloses a mechanism that uses
the
shape recovery of a shape memory alloy for sealing openings or high-pressure
passages.
A component made of a shape memory alloy can be processed in its martensitic
state to
have a reduced dimension smaller than that of the opening or the passage to be
sealed.
Upon heating, shape recovery takes place that is associated with the reverse
crystalline
phase transformation of martensite. The shape recovery of the previously
processed shape
memory alloy component yields a diameter greater than that of the opening or
passage to
be sealed. The shape recovery provides the dimensional interference and force
required
for sealing.
[0018] Wu, et al., builds on the work of Simpson, et al., U.S. Patent No.
4,770,725,
to use both the defined chemistry and the specified process method to effect a
specific heat
sealing application which employs a heat activated recovery transformation.
[0019] The wide thermal hysteresis available from thermal and mechanical
treatment
of alloys disclosed in the literature is attractive for articles which make
use of a thermally
induced configuration change, since it enables an article to be stored in the
deformed
configuration in the martensite phase, at the same temperature at which it
will then be in
use, in the austenite phase. This thermal and mechanical treatment is used in
a variety of
industrial heat-to recover couplings and connectors (L.Mcd. Schetky, The
Applications of
Constrained Recovery Shape Memory Devices for Connectors, Sealing and
Clamping,
Proceedings Super-elastic Technologies, Pacific Grove, Calif. (1994)).
[0020] It has been reported that a reverse transformation start temperature
As' has
been raised to +70 C after specimens were deformed to 16% strain at different
temperatures, where the initial states of the specimens were pure austenite
phase and/or
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martensite phase depending upon the pre-straining temperature regime. It was
found that
a transformation hysteresis width of 200 C could be attained and the reverse
transformation temperatures were measured by forcing a shape-memory recovery
via
heating, and that up to 50% of the pre-strain could be recovered. The work was
done by
Xiang-Ming He, et, al, Study of the Ni41.3Ti 38,7Nb20 Wide Transformation
Hysteresis
Shape Memory Alloy, Metallurgical and Materials Transactions, Vol. 35A, Sept
2004.
The work cited is an optimization of various pre-straining conditions to
maxiniize the
strain recovery possible for heat recovery application.
[0021] While the wide hysteresis confers certain advantages when the thermally
induced changes in configuration are to be exploited, a wide hysteresis in
stress-strain
behavior is generally inconsistent with the properties of an alloy that are
desirable in stent
or medical device applications.
[0022] Various methods have been described to deliver and implant stents. One
method frequently described for delivering a stent to a desired intraluminal
location
includes mounting the expandable stent on an expandable member, such as a
balloon,
provided on the distal end of an intravascular catheter, advancing the
catheter to the
desired location within the patient's body lumen, inflating the balloon on the
catheter to
expand the stent into a permanent expanded condition and then deflating the
balloon and
removing the catheter. One of the difficulties encountered using other stents
involved
maintaining the radial rigidity needed to hold open a body lumen while at the
same time
maintaining the longitudinal flexibility of the stent to facilitate its
delivery.
[00231 What has been needed and heretofore unavailable is a stent which has a
high
degree of flexibility so that it can be advanced through tortuous passageways,
can be
readily expanded, and yet have the mechanical strength to hold open the body
lumen into
which it expanded.
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[0024] Thus, it is desirable to develop an alloy that is very ductile and
uniquely suited
for deployment of medical devices, such as stents, into the human body. In
particular, for
stents, it is desirable that the compressed stent maintain its shape until
expanded.
SUMMARY
[0025] In one embodiment, the present disclosure provides a Ni-Ti ternary
alloy that is
particularly useful for medical instruments and devices, as well as components
thereof.
[0026] In another embodiment, the present disclosure provides an alloy having
variable hardness/stiffness properties and which is useful for medical
instruments and
devices, as well as components thereof.
[0027] In yet another embodiment, the present disclosure provides a material
for
making medical instruments and devices as well as components thereof that are
formable
without crack initiation sites when expanded or deformed.
[0028] Another embodiment is directed to a method of producing a Ni-Ti ternary
shape memory alloy, specifically strained induced austenite (SIA), with
improved
characteristics which are desirable for use in various applications such as
medical devices.
[0029] The disclosure provides for a stress induced martensite that is locked
in place
by the presence of a third element such that AS'>37 C . The material can be
placed inside
a ma.mmalian body that can have its As' -+ AS by controlled deformation, thus
creating a
material having mechanical properties superior to the mechanical properties
upon
insertion. When the material is used in a medical device structure, the device
can be
placed inside a mammalian body and restored to a stiffer mechanical state
after controlled
deformation or expansion.
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[0030] The material is particularly advantageous in that a medical device
comprising
the material can be placed inside a mammalian body that can have its AS'
initially reduced
to a temperature < 37 C by controlled deformation, and further restoration of
the
original AS and austenitic transformation of mechanical properties by heat
driven shape
memory recovery process. Once placed inside a ma.mmalian body, the material
can be
restored to a stiffer mechanical state after controlled deformation whereby
As' < 37 C,
and the austenitic transformation is completed by a combination of controlled
deformation
and assisted by temperature driven shape memory recovery process.
[0031] The disclosure is also directed to a method for the thermal mechanical
treatment of a Ni-Ti alloy performed at temperatures < Md (the temperature at
which
martensite can no longer be stress induced) where the hysteresis is widened
such that after
the treatment is completed AS < AS , or when used for medical applications, As
< 37 C <
As'.
[0032] The present disclosure provides a method of producing a Ni-Ti ternary
alloy
that exhibits properties desirable for medical instruments and devices. In
particular, the
treated alloy exhibits a high degree of ductility and low mechanical strength
properties
during insertion into the body and subsequently can be made stiff with
significantly higher
mechanical properties after being subjected to a controlled deformation of a
critical
magnitude.
[0033] The present disclosure is directed to a method for the thermal
mechanical
treatment of a Ni-Ti ternary alloy performed at temperatures < Md wherein the
hysteresis
is widened such that after the treatment is completed AS < As , comprising:
a) annealing or partially annealing a Ni-Ti ternary alloy through a high
temperature solution treatment followed by cooling;
b) mechanically straining the alloy be under a load while
simultaneously cooling the material to a temperature around or less than Ms,
thus
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shifting As up much higher such that AS < AS and retaining a sufficient amount
of
strain in the strained element whereby if controlled deformation is applied
the
alloy is transformed to the austenitic phase, shifting AS to A.
[0034] In preferred embodiments, the alloy is mechanically strained between
about 10
to about 25% under a constant load while simultaneously cooling the material
to a
temperature around or less than Ms, thus shifting As up much higher such that
AS < AS
and retaining about 2% to about 15% strain in the strained element whereby if
controlled
deformation is applied the alloy is transformed to the austenitic phase,
shifting AS to As.
[0035] Mechanical straining of the alloy during cooling forms martensite
variants
having volume fractions comprised mainly of (a) stress induced martensite and
or (b)
twinned and deformed martensite. The volume fractions formed are dependent
upon the
cooling rate and the applied pre-strain load stress. For instance if the
applied pre-strain is
at a temperature > Af and < Md, the microstructure may reveal formation of
stress
induced martensite and if the pre-strain is applied at a temperature < Af the
micro
structure may reveal formation of twinned and deformed martensite.
Optimization of the
martensite volume fractions are dependent upon the process pre-straining
conditions and
temperature. Reorientation of the martensite variants are prevented by the
presence of a
third insoluble element in the metal matrix giving rise to Af'.
[0036] Controlled deformation is displacement and the associated stress/strain
field
(bending, compression, tension, shear) required to initiate a complete or
partially
complete transformation wherein As' is shifted back to As. During controlled
deformation
the prepared material is deformed easily at low stress levels, and after a
sufficient range of
displacements have been completed, the material undergoes a permanent shift in
mechanical properties such that further displacement now occurs at higher
stress levels.
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[0037] By pre-straining the Ni-Ti ternary alloy using the method described
above, it
has been found possible to produce a shift of a normally austenitic alloy at
body
temperature, into an alloy that has martensitic properties at body
temperature.
[0038] Surprisingly, the controlled deformation causes the pre-strained
material to
become much stiffer where there appears to be almost a spontaneous conversion
from
martensite to austenite. The controlled deformation appears to "unlock" the
martensitic
alloy structure and cause the mechanical properties to revert seemingly
spontaneously
back to the austenitic alloy. Adding controlled deformation in the appropriate
amount
will cause the AS' to be restored to the original material A. In certain
embodiments of the
disclosure, the controlled deformation may occur through bending, compression,
tension
and/or shear stresses.
[0039] The resulting material is martensitic at all temperatures < As', very
ductile and
uniquely suited for deployment of medical devices such as stents into the
human body.
After pre-straining the material, the compressed stent maintains its shape
until balloon
expanded.
[0040] In one aspect of the disclosure, the shape memory alloys processed by
the
method comprise a terna.ry element E from 3 at.% to about 20 at.% of the alloy
composition E can comprises a ternary element that is effectively insoluble in
the Ni-Ti
matrix such as Nb, Ta or Zr, and the like.
[0041] The pre-straining process, in combination with the preferred Ni-Ti
alloy
composition employed renders the components flexible that, in turn, make
medical
components, in particular stents, easy to insert and deliver into the desired
location. The
resulting alloy is a material that can revert from martensite to austenite by
deformation
process as described above in (b). This characteristic is particularly
advantageous when
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trying to create a structure in the human body because delivery can be through
a small
opening and then after deployment the structure takes on the stiffer rigid
property.
BRIEF DESCRIPTION OF THE DRAWINGS
[0042] The foregoing and other objects and features of the disclosure will
become
apparent from the following description of preferred embodiments of the
disclosure with
reference to the accompanying drawings, in which:
Figure 1 is a graph showing radial expansion force versus stent diameter for a
316L or Co-Cr annealed stent material;
Figure 2 is a graph of temperature vs. strain showing the shift of A, -+AS'
after an
alloy material [ 1] has been processed by the first part of method while [2]
is an example
of an alloy according to the present disclosure, processed by the method;
Figures 3a and 3b shows two graphs comparing pre-strained Ni-Ti-E alloy
heating
(I) and deformation (II) recovery modes;
Figure 4 is a schematic representation of an apparatus used to perform a three
point bend test to show the onset of "unlocking" mechanism;
Figure 5 is a graph of a three point bend test performed with the apparatus of
Fig.
4. These plots show first and second bending cycles of strain induced
austenite and the
unloading and reloading which occur at 0.060 thousandths (")(0.1524 centimeter
(cm)
displacement giving some indication of the spring back characteristic;
Figure 6 is a schematic representation of some components used in the Acculine
bend moment tester used to provide cyclic bend testing with a constant moment
arm;
Figure 7 is a graph of the test results of the tests run with the apparatus of
Fig. 6
showing strain induced austenite undergoing stiffness transformation during
its first and
second bending cycle;
Figure 8 is an apparatus to measure of the generic stent cell performance
characteristic;
Figure 9 is a graph of data of the test using the apparatus of Fig. 8;
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Figures 10A - 10C show an apparatus used for a test that shows the material
transformation occurring by cyclic bending of the Strain Induced Austenite
(SIA)
material;
Figure 11 is graph of the data using the apparatus in Figures 10A - 10C;
Figure 12 is a graph showing stent diameter vs. radial expansion of a stent
prepared from the strain induced austenite; and
Figure 13 shows the strain behavior for a stent fabricated from strain induced
austenite in relation to Figure 11.
DETAILED DESCRIPTION OF THE ILLUSTRATIVE EMBODIMENTS
100431 Most stents on the market today are made from materials such as 316L or
Co-
Cr (L605 and MP35N) materials. The stents are compressed, permanently deformed
and
crimped onto a delivery catheter equipped with an underlying balloon. Once
positioned in
the lumen, a balloon is inflated to outwardly expand the stent material beyond
its yield
point, until the stent makes contact with the vessel wall.
[0044] Figure 1 shows how the radial expansion force varies as a function of
stent
diameter. These stent materials have four limitations: (1) the expansion force
increases
with expansion diameter (Figure 1[ 1]); (2) material spring back requires
(Figure 1 [2])
over expansion to achieve the final diameter and this results in vessel
trauma; (3) the high
force to expand the stent requires an inflation balloon to be of a heavy wall
thickness
increasing the overall profile and stiffness of the delivery system; and (4)
under X-Ray
fluoroscopy some of the above mentioned materials have poor visibility.
[0045] Additionally, some stents on the market today are made from
superelastic
NITINOL. In these devices, the "active" Af (as defined by ASTM F2082) of the
final stent
structure is between 0 C and 37 C, more particularly between 10 C and 25 C.
These
stents are radially compressed at (a) room temperature against a radial
outward force
generated by the upper plateau stress (about 70 ksi or 483 MPa) of
superelastic NiTttvor.
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or (b) at a temperature below MS against the mechanical stress (about 12 ksi
or 83 MPa)
of martensite. In either case, at room temperature the outward radial force of
the
compressed stent diameter can be constrained with a delivery sheath fitting
over the
compressed stent. These stent materials have the following limitations: (1)
the
constrained outward force of the stent requires an outer sheath which adds
unwanted
stiffness and bulk to the delivery system; (2) the continually outward force
of the
compressed stent exerts a frictional force on the sheath which increases the
sheath
retraction force; (3) once the sheath is pulled back the stent rapidly expands
to the final
diameter which may cause vessel trauma; (4) the stent diameter cannot be over
expanded
to acconunodate any gap between the vessel lumen and the selected stent; and
(5) under
X-Ray fluoroscopy, binary NITINOI. has poor visibility.
[0046] In one aspect of the disclosure, the shape memory alloys processed by
the
method comprise a ternary element E from 3 at.% to about 20 at.% of the alloy
composition, wherein E can comprise a ternary element that is effectively
insoluble in the
Ni-Ti matrix. Examples of suitable ternary elements include, but are not
limited to,
niobium (Nb), tantalum (Ta), Zirconium (Zr) and the like. It is also thought
that shape
memory alloys comprising combinations of element E can be effective.
TABLE I
Com osition of Materials Suitable for "Strain Induced Austenite"
Alloy Composition Titanium Atoniic %) Nickel Atomic % E (Atomic %)
Ni-Ti-E Y% Z% 0.1 <X%<20%
1. X is the percentage addition of a third insoluble element (E) such as
Niobium (Nb),
Tantalum (Ta) and other insoluble ternary elements into the Nickel Titanium
matrix. The
element is supplied in sufficient amounts to optimize the "locking and
unlocking"
mechanism of "strain induced austenite."
2. Y and Z are used to adjust the ratio of Nickel and Titanium necessary to
achieve the target
ingot A temperature considered optimal by the end application. The ratio of
Atomic Percent
Ni to Atomic Percent Ti is from about 1.6 to .60.
100471 In particular, the insoluble ternary element (Nb, Ta, or Zr) appears to
play a
role to "lock" the As' structure during the pre-straining cycle in such a way
that
restoration of the austenite state is prevented until either (a) the materials
is heated above
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AS and shape recovery takes place or (b) the controlled deformation sets up a
strain field
sufficient to trigger material "unlocking" and restoration of original
material.
[0048] In contrast with other materials, the alloy of present disclosure is
particularly
well suited for structures that are packed in a small configuration, easily
expanded with
rniniunal force, and later after deployment becomes significantly stiffer, a
property that is
desired for various medical components, particularly stents.
Example 1 - Preparation of the pre-strained alloy
[0049] The method was applied to Ni44T47Nb9 alloy to provide the response
illustrated in Figure 2 which showed an increase of As --> As' and the
widening of the
hysteresis from [1] to [2]. The alloy was partially or fully annealed between
450 C to
900 C, particularly 650 C to 900 C for <60 minutes wherein after cooling the
austenite
start temperature As was about -65 C. The material was strained under load
while the
material was cooled, preferably to a temperature between the martensitic start
temperature of MS about -85 C and the As about -65 C. The material was then
strained
8% to 25%. Upon completion of the pre-straining cycle AS was elevated above 37
C and
particularly above 60 C.
[0050] Another aspect of the disclosure is the effective insolubility of
niobium in the
Ni-Ti matrix. It appears that during the treatment method, there is a
partitioning of the
total strain of the material and each component (i.e. nickel-titanium and beta
niobium)
seems to work as an individual component at similar levels of flow stress.
Depending
upon the pre-straining temperature regime, martensitic Ni-Ti deforms
reversibly by either
twin boundary motion or stress induced martensite motion, while the beta-
niobium
deforms irreversibly via slip to "lock" the martensitic deformations in place.
In either case
the locked structure mechanism corresponds to the new AS .
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[0051] After treatment, the Ni-Ti-Nb microstructure is comprised of: (a)
stress
induced twined martensite consisting of type I(111) M twins and occasionally
(001) M
twins and anti-phase domain boundaries (Icomet-92); and (b) the sofft beta-
niobium
particles are deformed irreversibly via slip.
[0052] The plentiful block shaped niobium phases are adjacent to the different
martensite variant boundary below M,. After pre-straining, these block phases
serve as
obstacle sources to inhibit the reorientation of the martensite, so martensite
reorientation
has difficulty proceeding at lower stress levels. When stress is added by
controlled
deformation (e.g., bending), the martensite reorientation is free to proceed
once the
blocking constraint of the niobium phases is removed by the additional stress.
[0053] Another aspect of the method is the creation of expandable structures
that can
be collapsed and ready for small opening insertion, that do not required
external
constraints. In this way, the pre-strained structure is comprised of (a)
stress induced
martensite or (b) twinned and deformed martensite which is "locked" in place
by the
present of the third and insoluble element (in this case the soft niobium
particles). This
type of structural element may not need an external cover to hold back the
restoring force
typically found in superelastic NrrNoL expandable structures. One advantage to
the
method is that no external cover is required for insertion, leading to a
reduced design
profile that is advantageous for a number of medical device applications.
[00541 A tubular stent element can be manufactured by one of the following
series
(e.g., 1, 2 or 3) of process steps. Note that the thermal mechanical treatment
can occur
either before or after the machining step in items 1 and 2 below.
1) Starting Tube - Laser or Chemical Machining - Pre-Straining to Increase
AS - Deployment and Expansion
2) Starting Tube - Pre-Straining to Increase AS Laser or Chemical
Machining - Deployment and Expansion
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3) Starting Wire - Shape Setting - Pre-straining to Increase AS' -
Deployment and Expansion
[0055] Other appropriate manufacturing methods incorporating the treatment
method
would be known to those skilled in the art for use in applications with wire,
tube, strip or
appropriate forms.
[0056] As discussed above, the present disclosure offers an alternate method
to
deploy structures and devices in the human body that are made from other
materials (i.e.
316L and Co-Cr alloy systems) and super elastic NiTINOL.
[0057] When the pre-strained Ni44T47Nb9 material is heated such that T> AS',
the Ni-
Ti matrix reverts to its parent phase (Table 1- Process I) with a
corresponding increase
of mechanical properties (Figure 3 - Process I) sufficient to overcome the
"locking" force
of the deformed beta niobium particles and thereby "unlocking" the structure
and
recovering both the original pre-strained shape and austenite start
temperature (AS). The
drawback to the resulting material is that heating is not a preferred method
to deploy a
medical device into a mammalian body.
Table I- Illustrates the Process Differences Between Process I (Heating) and
Process II(Controlled
Deformation) which is Strain Induced Austenite for Ni44 Ti47 Nb9.
Process Initial Pre-Straining Application Recovery Phase
Method Preparation
I Cold Working Pre-strain while Constrained Heating T> AS'
plus Annealing cooling article to Recovery Application To achieve: As'-. A,
@ about Ms or < Ms (e.g. pipe coupling) and shape recovery
450 to 850 C temperature to
1 to 30 minutes achieve: .AS- >A~'
II Cold Working Pre-strain while Biomedical Device Controlled deformation
plus Annealing cooling article to Insertion (e.g. stent) of sufficient scale
to
450 to 850 C about Ms or < Ms achieve: As'-. A
1 to 30 minutes temperature to
achieve: AS--)AS'
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[0058] In the present method, the pre-strained material is "unlocked" by a
controlled
deformation (Table 1- Process II) and upon reaching a critical strain level
(c,) (Figure 3
- Process II), the soft martensitic structure begins to revert to the parent
phase having
improved mechanical properties useful in many medical devices including
stents.
[0059] Specifically, a controlled deformation of sufficient strain (cc) can
cause the
parent phase to preferentially nucleate in a jump-like manner. At locations
where flow
stress levels are sufficiently large, soft niobium particles no longer
constrain martensite
variants which are now "unlocked" an undergo a parent phase transformation as
a series
of stress level reductions (Figure 3 - Process II).
Example 2 - Three point bend test
[0060] Referring to Figs. 4 and 5, a three point bend test was performed to
demonstrate the "unlocking" clicking phenomena, the spring back characteristic
in
the "unlocking region" at 0.060" (0.15 cm) displacement, and to show the
stiffness transformation between the first and second bend cycle.
[0061] A 0.020" (.05 cm) 316LVM (low carbon vacuum melted) wire in the
fully annealed condition (typical of condition and chemistry for stent
material) was
prepared.
[0062] For the Strain Induced Austenite (SIA), 0.020" (.05 cm) Ni-Ti-Nb was
pre-strained under approximately 302 ksi (2082 MPa) during a cooling cycle
from
20 C to -140 C. The wire had about 10% retained strain after returning to room
temperature.
[0063] A test apparatus 10 for the above method is shown in Figure 4 below.
Wire specimen 12 was placed across a 0.375" (.952 cm) unsupported span 14. An
Instron 5544 equipped with a 20 pound (lb.) (9.072 kilogram (kg)) load cell
and
Blue Hill software was used for data capture. Anvil 16 was secured in the
Instron
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cross head and (a) advanced down at .04"/minute (0.10 cm/min) for a distance
of
0.060" (0.15 cm) (b) moved upward until the load reached zero, and (c)
advanced
down until total displacement equaled 0.100" (0.254 cm).
[0064] Figure 5 shows the load displacement behavior for 0.020" (0.05 cm)
SIA wire 12, subsequent bending cycle after a strain induced austenitic
transformation and a 316LVM fully annealed wire of equivalent cross section.
The "unlocking mechanism" of the SIA material 12 was apparent in the
displacement range from about 0.03" (.076 cm) to 0.08" (.2032 cm). The onset
of
material "unlocking" occurred at 0.03" (.076 cm) and the load capacity was
suddenly shifted from 0.281bs (127 grams (g)) to 0.161bs (72 g). There were a
number of other small drops in load capacity observed up to 0.080" (0.203 cm)
of
displacement. The net effect of the unlocking mechanism is that bend force for
the SIA material did not grow very large. In comparison, the performance of
316L fully annealed material (with well documented extremely low stiffness and
ductile material properties) was perhaps stiffer by a factor of two throughout
the
range of displacement.
[0065] At 0.060" (0.15 cm) wire 12 was unloaded to zero and then the load
was reapplied. The spring back for the SIA material was about 0.015" (.038 cm)
and the 316LVM fully annealed material was about 0.009" (0.0229 cm). The test
was stopped at 0.93" (2.36 cm) of displacement. The peak force for SIA was
0.33 lbs (0.15 kg) while the peak force for 316LVM was 0.62 lbs (0.28 kg). The
data demonstrates that: (a) the initial bending stiffness of SIA is about one-
half
(1/2) that of annealed 316 LVM, (b) the "unlocking" phenomena is prevalent in
the
range between 0.03" (0.0762 cm) and 0.08" (0.203 cm) of displacement and is
responsible for keeping the stiffness extremely low during initial bending. In
addition, it is well documented that fully annealed 316LVM has yield strength
of
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approximately 45,000 psi (310 MPa), whereas the first bend cycle of SIA
material
has a yield point considerably less than the 316LVM.
[0066] After the initial bending cycle the deformed SIA wire 12 was
straightened and retested under the same 3 point bend test conditions. These
results are shown in Figure 5 and identified as SIA - Subsequent Bend Cycle.
At
about 0.06" of deflection the subsequent bending cycle after strain induced
austenitic transformation showed a wire 12 that was more than twice as stiff
as the
0.020" 316LVM fully annealed wire and more than 4 times stiffer than the
0.020"
SIA wire during the first bending cycle. This is evidence of the bending
induced
transformation from martensite to austenite.
Example 3 - Observation of the Reverse Phase Transformation by Simple Bending
[0067] Further evidence of strain induced austenite was easily observed using
wires having a starting diameter of 0.020" which, after the pre-straining
sequence
described above, resulted in stable martensite material having a diameter of
0.0192". As described above, there are two paths by which the original 0.020"
wire diameter can be recovered: (1) by simple bending and (2) by heating. The
heat to recovery method for these alloys has been successfully employed for
industrial applications but is not considered acceptable for medical
applications
because of the high temperature (e.g., greater than 60 C) required to effect
the
transformation. However, if the pre-strained wire is bent between ones fingers
at
room temperature, the transformation taking place can be observed by a soft
clicking action that ends in a stiff bent wire. If the bent wire is now
straightened
by bending in the opposite direction more soft clicking can be observed. The
end
result is a phase transformation of wire from a soft martensitic material into
a
stiffer austenitic material. Proof that a reverse transformation has taken
place can
be found by measuring wire diameter with calipers. After bending the pre-
strained
wire as described above, wire recovered in diameter from 0.0192" after pre-
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straining, to its original 0.020" diameter. This growth in wire diameter is
proof of
the recovery transformation via bending.
Example 4- Acculine Bend Moment Testing
[0068] Referring to Figures 6 and 7, this test compared the bending moment
transformation behavior of 0.020" pre-strained Ni-Ti-Nb as described above and
0.020" superleastic straight binary NITINoL. The bending stiffness of strain
induced austenite and superelastic NITrtvoL straight wire samples were tested
under identical conditions using an Acculine AE#-BM Bend Tester 110. Each
0.020" wire specimen 112 was mounted between custom mounting blocks 114
with a 1 millimeter (mm) bend radius. Drive pins 116 and 118 rotated at 9.0
degrees per second ( /sec) to maintain a constant 0.3 cm moment arm while wire
112 was deflected from the vertical position (0 ) counter clockwise to + 45 ,
returned to 0 , counter clockwise to -45 and back to 0 to complete one bend
cycle. Data was collected by a 10 in-oz rotary torque measurement sensors 120
and 122 and captured for graphical Microsoft Excel presentation. Figure 6
shows
the experimental set up. Figure 7 is a graph of the data from the first,
second and
third bend cycle. The first bend cycle shows a peak moment for the Ni-Ti-Nb to
be about 41 % less then superelastic NinvoL, the second bend cycle shows the
peak maximum bend moment to be about 13% of the peak moment of superelastic
NinNoL, and the third cycle shows the Ni-Ti-Nb is approaching the performance
of the superelastic NiTTNOL wire. This is further evidence that pre-strained
Ni-Ti-
Nb had undergone a transformation from a locked Ni'ritvoL structure (Af') to
an
unlocked structure restoring the original Af.
Example 5 -- Expansion of Two Parallel Wires
[0069] Referring to Figs. 8 and 9, the behavior of the material during
expansion was
tested, specifically, 0.020" (.05 cm) 316LVM (low carbon vacuum melted) wire
in the
fully annealed condition (typical of condition and chemistry for stent
material). SIA
0.020" (0.05 cm) Ni-Ti-Nib pre-strained under approximately 302 ksi (2082 MPa)
during
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a cooling cycle from 20 C to -140 C. The wire had about 10% retained strain
after
returning to room temperature.
[0070] Referring to Fig. 8, the testing apparatus 210 comprised two small
Delran blocks 212 and 214 which were devised to hold two 0.020" (0.05 cm)
wires 216
and 218 in a parallel starting configuration (wire spacing 0.100" (0.254 cm)).
The
spacing 220 between blocks 212 and 214 used to clamp the parallel wires 216
and 218
was about 1.414" (3.59 cm). An Instron 5544 equipped with a 201b (9.072 kg)
load cell
and Blue Hill software was used for data capture. Specially designed micro-
lifting hooks
222 and 224 (0.080" or 0.203cm width) were designed to expand the parallel
wires by
hooking to the mid point of the 1.414" (3.59 cm) span and then traveling a
total distance
of 1" (2.54 cm) at a rate of .25 inch per minute ("/min) (.635 centimeter per
minute
(cm/min) while recording the load. Once the crosshead reached a travel
distance of 1"
(2.54 cm), the expanded structure was unloaded until the crosshead reached a
travel
distance of.96" (2.44 cm).
[0071] As shown in Fig. 9, the load deflection behavior of 0.020" (0.05 cm) Ni-
Ti-Ni
wire has compared with the 316LVM fully annealed wire of equivalent cross
section. The
"unlocking mechanism" of the Strain Induced Austenite (SIA) material was
plainly visible
(fine jagged pattern) in the extension range from about 1" (0.254 cm) to 0.7"
(1.8 cm). In
this range there was negligible force increase. The resistance to bending of
the 0.020"
(0.05 cm) SIA was less than half ('/z) that of the 0.020" (0.05 cm) 316LVM
fully annealed
material. It is quite well known that 316LVM in the fully annealed condition
has a yield
strength of approximately 45,000 psi (310 MPa) and this would lead us to state
that the
yield point of the SIA material on the first cycle was considerably less than
316LVM,
which is already an extremely soft ductile material. This curve also shows
that the work
required to achieve a given displacement was considerably reduced to achieve a
given
displacement when compared with 316LVM fully annealed material. After
expansion to
1" (2.54 cm), the curves were unloaded to 0.96" (2.43 cm) and the load was
recorded.
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During unloading the slope of the spring back was calculated and it was found
that the
0.020" (0.05 cm) SIA material was 9.058 pounds per inch (lb/in) (1.411 newtons
per
millimeter (N/mm) while the 316LVM fully annealed material was 16.61bs/in
(2.907 N/mm).
Example 6 -- Material transformation through bending of the SIA material.
[0072] Referring to Figures 10A-10C and 11, a test was constructed to show the
material transformation occurring by bending of the SIA material.
[0073] SIA -0.020" (.05 cm) Ni-Ti-Nb pre-strained under approximately 302 ksi
(2082 MPa) during a cooling cycle from 20 C to -140 C. The wire had about 10%
retained strain after returning to room temperature.
[0074] A straight length of 0.020" (0.05 cm) SIA material specimen 310 was
clamped
in mounting blocks 320 at one end and at a distance of 0.20" (0.5 cm) from the
fixed end,
a close fitting guide 330 machined from a plate (0.080" or .2 cm width) was
attached. The
close fitting guide 330 was attached to the crosshead 340 of an Instron 5544
equipped
with a 201b (9.072 kg) load cell and Blue Hill software for data capture.
Using the
crosshead extension, the test began at the neutral position "A" (zero
deflection with zero
load). The cross head cycled upward +0.1" (-0.25 cm) to position "B" and
downwards -
0.2" (-0.51 cm) to Position "C" and then upward +0.2" (+0.51 cm) to Position
"B" and
so forth until the test was completed. The cross head speed was 0.2"/min or
about 0.5
cm/min.
100751 As shown in Figure 11, start at (A) and progress along with low force
with
evidence of "unlocking" to (Point B). The material was bent in the reverse
direction from
(Pt B) to (Pt A) with no force (more unlocking) until it got back slightly
past the origin at
(Pt A). Continued bending towards (Pt C) caused rapidly increasing stiffness
with the
force at (Pt C) more than twice the force at (Pt B). During the second round
trip to Pt
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B 1 and Pt C l, the results confirmed the material transforma.tion was
completed,
demonstrating that mechanical bending converted AS to As for this material
system (i.e.,
Process II of Table I). The details are discussed below.
[0076] The results in Figure 11 show the load deflection behavior of 0.020"
(0.19 cm)
SIA wire for negative and positive bending conditions.
Displacement A-+B
[0077] The test began at Position A (zero deflection and load). During
positive
bending to Position B, there was evidence of the "unlocking mechanism" at
about 0.02"
(0.5 cm) when the load was reduced to near zero level (shown by a large
instantaneous
spike). The load recovered and overall stiffness remained relatively low as
the
displacement increased to about 0.07" (0.17 cm) at which point the stiffness
increased
quickly.
Displacement B --> C
[0078] During unloading, the load reduced quickly to a displacement of 0.075"
(0. 19 cm) as would be expected. Beyond this range, there was evidence of
additional
material "unlocking" as the specimen bent easily without resistance and with
near zero
stiffness in the displacement range of 0.07" (0.18 cm) -> -0.01" (0.0254 cm).
At the
displacement of -0.01" (.0254 cm), the "unlocking mechanism" appeared
complete, and
the specimen has now transformed into a much stiffer material as evidence by
the
increasing load at displacement Position C.
Displacement C --+ B 1
[0079] The material has been completely transformed and the load at B 1 has
increased by more than twice when compared with Position B. Additionally, in
the
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displacement range from 0.0" (0 cm) ---- 0.05" (0.127 cm), the load during the
second
cycle was four (4) to six (6) times greater than on the first cycle. This
material test shows
that a stiffness transformation has occurred and the "unlocking phenomena"
observed in
the first displacement cycle has disappeared.
Displacement B 1--+ C 1-->=A
100801 Subsequent cycling shows the material is no longer exhibiting the
"unlocking"
phenomena and bent with considerably more stiffness than during the first
cycle.
100811 This type of curve certainly demonstrates that mechanical bending
converted
AS to A, for this material system (i.e., Process II of Table I). The near zero
stifffness
during bending, followed by a transformation to a significantly stiffer wire
is unique, and
offers the potential to engineer many useful devices.
[0082] A further advantage of the present disclosure is to provide a thermal
mechanical treatment regime such that AS' is sufficiently > 37 C, such that
little or no heat
induced shape memory recovery occurs during temperature exposures caused by
placement of the stent into the human body, curing of drug coatings,
sterilization or
shipping of finished medical devices.
[0083] In another aspect of the disclosure, it is possible to deploy a highly
elastic
(super-elastic) stent without having to constrain and de-constrain stress
induced
martensitic stent structure in contrast to, for example, Jervis, U.S. Patent
No. 5,067,957,
discussed above. In this aspect of the disclosure, a Ni-Ti-E material (as
described above)
having appropriate chemistry to yield a fully annealed temperature AS about -
15 C and
ideally processed to have super-elastic properties at body temperature
(particularly having
an "active" Af between 10 C and 20 C) can be "locked" by the thermal
mechanical
treatment means such that AS > 37 C. When placed in the human body without
constraint and subsequently balloon expanded, the structure can "unlock" at
which point
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A{ again approaches Af. Applying the straining techniques above, such a
material's As
could be shifted higher and the resulting structure would be martensite at
body
temperature. At this point, insertion into the body followed by the controlled
deformation
would be sufficient to restore the original As and therefore the superelastic
properties at
room temperature. This alternate deployment method eliminates the force
required to
hold the stent in the collapsed state, reduces high friction loads between the
stent and
delivery sheath and minimizes the stent delivery profile.
[0084] By reverting the material back to austenite (as AS'--+ As), the
superelastic Ni-
Ti-E properties are reached at a certain point and then it can be further
expanded easily
when needed - a property particularly desirable for stent material.
Furthermore, the
material would possess the advantages of radiopacity and higher strength when
compared
to binary NITINOL.
[0085] A stent fabricated from "strain induced austenite" as described in this
disclosure can have a much different stent radial deployment force than those
of other
alloys or superelastic NITItvOL. Referring to Figure 12, initially the stent
radial expansion
force increases in a linear reversible fashion with stent diameter and
corresponds to the
elastic deformation limit of martensite along the path from the origin --+ P.
The radial
force required to achieve point P is much less than what is required to
achieve the yield
point of other materials (316L and Co-Cr). This is a particular advantage of
the current
disclosure. The "unlocking" process of martensite begins at point P and ends
at point B.
As the stent diameter increases, the required radial outward force is constant
or
decreasing during the "unlocking" process P--+B where the material completes a
reverse
transformation to austenite.
[0086] Advantageously, when the diameter of the "unlocked" stent is sized
appropriately to the vessel lumen (this would correspond to point B in Figure
12), then no
overexpansion of the lumen is required, and thus vessel trauma is el'uninated
or
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minimized. Furthermore if the vessel wall tries to contract or collapse, the
stent offers a
reserve of radial resistive force, as shown Figure 13, IV. The reserve of
radial resistance
is coming from the material transformation of martensite to austenite by the
controlled
deformation of the stent expansion cycle. The change in material stiffness is
shown by
examining Figure 5 SIA - subsequent bending cycle.
[0087] In the present disclosure, the force required to reach the fully
deployed stent
diameter (point B Figure 12) can be much less than the expansion force
required by a
stent made from other materials. The reduced expansion force can lead to an
optimization of the stent delivery profile.
Biomedical Device ANlications
[0088] In general, systems using the present material can provide higher
mechanical
properties than other binary alloys (for example, NiTiNOL), resulting in
smaller device
cross sections and minimal design profile. Such devices can reduce trauma
since they do
not have to be overdeformed during deployment, as in the case of materials
such as
stainless steel (316L), Co-Cr alloys (L605, MP35N), and titanium-based
materials.
Another advantage of the present materials is less inflation pressure of the
balloon. The
addition of Nb or Ta into the Ni-Ti-E alloy can improve the radio-opaque
properties of
the material, allowing doctors to find the location of smaller cross sections
under X-Ray
fluoroscopy. The alloy exhibits nonmagnetic, low torque properties, and offers
a crisp
image under MRI imaging which is a medically desirable property.
[0089] For percutaneous, intraluminal and laproscopic medical device
applications,
the present disclosure offers multiple advantages including: very low
deployment forces,
delivery systems with more flexible and smaller cross sections, and inflation
balloons with
thinner cross sections and lower operating pressures for safer and higher
reliability.
Designs using the present material do not have to be held in the compressed
position
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awaiting deployment, such as binary NITINor. at high stress levels (>60 ksi or
413 MPa)
during shipping, sterilization and storage.
Stents and Stent Grafts
[0090] As discussed above, stents are fabricated from laser cut tubes,
braided, coiled
or formed wires fabricated into tubular structures and used to repair the
patency of
narrowed, previously weakened or ballooned and otherwise impaired lumen or
other body
channels. They are deployed by the use of catheters in percutaneous,
intraluminal or
laproscopic procedures. Examples are: blood vessels, bile duct, esophagus,
urethra,
trachea and the like. Specifically: carotid and coronary vessel, interluminal
lining of aortic
abdominal aneurysms, iliac or femoral aneurysms, recanalization of injured
vessels caused
by blunt or penetrating trauma, dilation and recanalization of stenotic
arterial segments,
tampanade and obliteration of esophageal varices, recanalization of esophageal
stenoses
secondary to carcinoma or benign strictures, ureteral strictures and tracheal
strictures. In
all these applications, the present shape memory alloy would be advantageous
in its ease
of deployment.
[0091] The present disclosure improves on the current state of the art in
several ways
and the specific advantages depend upon the base material system in the
comparison.
[00921 For example, the present disclosure improves over a Nrrnvol, stent by
having
little or no outward radial force when placed in the delivery system tube. A
binary
NITINOI, stent exerts a chronic outward force on the inside waii of the
delivery system,
and during storage the inside wall of the delivery system sheath may become
imprinted by
the stent frame. During deployment of the Nirllvol, stent, the frictional
forces may be
quite high, whereas devices formed from the present alloys deploy more easily
and
provide a more flexible and reduced delivery sheath cross-section.
Furthermore, after the
expansion of a stent made from the present alloys is expanded, and the
transformation
described herein is complete, the new material can provide stiffer
characteristics than, for
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example, binary alloys such as NiTri'r0L. A further advantage is that niobium
is very
radiopaque under x-ray fluoroscopy whereas NiTtNOL is not.
[0093] When compared with a 316L stent, the present disclosure reduces the
delivery
system profile, the balloon expansion pressure is reduced and the total amount
of work
required to deploy the stent system is also reduced. The expansion force of a
316L stent
increases linearly as the stent diameter is expanded while the present
disclosure can
achieve significant stent expansion at near zero force as the diameter
expands. The lower
expansion force characteristic leads to a reduced cross section of the balloon
inflation
catheter, thereby leading to a lower profile and improved flexibility for the
delivery
system. A further advantage is that niobium is very radiopaque under X-ray
fluoroscopy,
whereas 316L is considered to have poor visibility.
100941 Figure 13 is an example of how a stent fabricated from strain induced
austenite
works overlayed with the data from Figure 11. The displacement axis represents
a small
about of bending taking place in the struts of individual stent cells. These
tiny stent cell
displacements could also represent changes in stent diameter. Roman Numerals I
thru IV
show the stent deployment relative to the material properties. Therefore, I
shows a
prepared stent loaded into the delivery system, II shows the near zero
deployment and
expansion force, III shows the fully expanded stent at the location of
complete material
transformation, and IV shows how the transformed stent material can now
provide
resistance to chronic radial force.
Sur ic~ al Staples
[0095] Surgical staples are typically made from 316L and titanium wire having
been
formed and loaded into delivery magazines. In many applications, a delivery
magazine
can hold 100 or more staples and they can be simultaneously fired at once. It
is desirable
to lower the combined firing force to push multiple staples from the magazine
holder and
simultaneously crimp the wire staples into the traditional B-shape profile.
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CA 02603738 2007-09-25
305349.3000-100
-30-
100961 The present application can improve upon multiple staple firing systems
by
reducing the total work and maximum force required to deploy a given number of
staples
and compress those staples into the required B-shaped profile. The initial low
stiffness of
the present disclosure allows engineers to redesign either: (a) surgical
staple devices that
are easier for the physician to grasp and fire, or (b) allow engineers to
design surgical
staple guns that fire more staples for an equivalent grasping force.
Medical Expansion Bolt/Bulkhead Connector Applications
[0097] It is also possible to utilize the strain induced austenite disclosure
to design
and fabricate a device that can fit through a small blind hole. Once inserted
and
mechanically expanded, the device can take on the new structural shape of a
particular
design intent that can prevent its removal. One particular application is the
Vascular Hole
Closure device in which a vessel wall is sealed from the outside wall. This
can be
achieved by inserting a device comprising the material into the vessel hole
(resulting from
a previous medical procedure) and causing a controlled deformation deployment
means
expanding the device in such a way that it cannot be removed, and thereby
seals the hole.
A second similar application in which the material may be useful is the atrial
septal defect
device.
[0098] It will now be apparent to those skilled in the art that other
embodiments,
improvements, details and uses can be made consistent with the letter and
spirit of the
foregoing disclosure and within the scope of this patent, which is limited
only by the
following claims, construed in accordance with the patent law, including the
doctrine of
equivalents.
What is claimed is:
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