U.S. patent application number 15/399632 was filed with the patent office on 2017-04-27 for bioabsorbable stent that modulates plaque geometric morphology and chemical composition.
The applicant listed for this patent is Abbott Cardiovascular Systems Inc.. Invention is credited to Syed F.A. Hossainy, James P. Oberhauser, Richard Rapoza, Yunbing Wang.
Application Number | 20170112646 15/399632 |
Document ID | / |
Family ID | 42398362 |
Filed Date | 2017-04-27 |
United States Patent
Application |
20170112646 |
Kind Code |
A1 |
Rapoza; Richard ; et
al. |
April 27, 2017 |
BIOABSORBABLE STENT THAT MODULATES PLAQUE GEOMETRIC MORPHOLOGY AND
CHEMICAL COMPOSITION
Abstract
Methods of treating a diseased blood vessel exhibiting stenosis
with a bioabsorable stent are disclosed. The implanted stent
supports the section of the vessel at an increased diameter for a
period of time to allow the vessel to heal. The stent loses radial
strength sufficient to support the section of the vessel in less
than 6 months after implantation, loses mechanical integrity, and
then erodes away from the section. The biodegradable stent results
in changes in properties of plaque with time as the stent degrades.
The time-dependent properties include the luminal area of the
plaque and plaque geometric morphology parameters.
Inventors: |
Rapoza; Richard; (San
Francisco, CA) ; Wang; Yunbing; (Sunnyvale, CA)
; Oberhauser; James P.; (Saratoga, CA) ; Hossainy;
Syed F.A.; (Hayward, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Abbott Cardiovascular Systems Inc. |
Santa Clara |
CA |
US |
|
|
Family ID: |
42398362 |
Appl. No.: |
15/399632 |
Filed: |
January 5, 2017 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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12510076 |
Jul 27, 2009 |
9572692 |
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15399632 |
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12364321 |
Feb 2, 2009 |
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12510076 |
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12506121 |
Jul 20, 2009 |
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12364321 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61F 2002/91566
20130101; A61F 2/06 20130101; A61L 31/06 20130101; A61L 2300/412
20130101; A61F 2230/0054 20130101; A61L 31/06 20130101; A61F
2250/0067 20130101; A61F 2210/0076 20130101; A61F 2/82 20130101;
A61F 2002/825 20130101; A61F 2/915 20130101; A61F 2/958 20130101;
A61L 31/16 20130101; A61F 2210/0004 20130101; A61F 2/91 20130101;
A61L 31/148 20130101; C08L 67/04 20130101; A61F 2002/91583
20130101 |
International
Class: |
A61F 2/915 20060101
A61F002/915; A61L 31/14 20060101 A61L031/14; A61L 31/16 20060101
A61L031/16; A61L 31/06 20060101 A61L031/06 |
Claims
1. (canceled)
2. A method of treating a diseased section of a blood vessel,
comprising: implanting a bioabsorbable polymeric stent comprising a
scaffolding composed of a pattern of struts at a diseased section
of a blood vessel to form a stented segment comprising the stent
and a vessel wall at the diseased section, wherein implanting
comprises deploying the stent from a crimped diameter to a diameter
of 3 mm, wherein a reference vessel diameter is 3 mm and a diameter
stenosis of the diseased section of the blood vessel
pre-implantation is greater than 50%, wherein the vessel wall
comprises a plaque region including a necrotic core component and a
fibrous component, wherein at pre-implantation and immediately
following implantation of the stent the necrotic core component is
in contact the lumen.
3. The method of claim 2, wherein as the stent degrades the fibrous
component becomes positioned between the necrotic core component
and the lumen so that the necrotic core component is no longer in
contact with the lumen.
4. The method of claim 2, wherein the scaffolding is made from
PLLA.
5. The method of claim 2, wherein the scaffolding polymer has a
degree of crystallinity between 25-50%, a Tg between 10 and
30.degree. C. above human body temperature, induced circumferential
polymer chain orientation to provide the sufficient radial strength
and to inhibit failure of the scaffolding as it supports the vessel
wall.
6. A method of treating a diseased section of a blood vessel,
comprising: implanting a bioabsorbable polymeric stent comprising a
scaffolding composed of a pattern of struts at a diseased section
of a blood vessel to form a stented segment comprising the stent
and a vessel wall at the diseased section, wherein implanting
comprises deploying the stent from a crimped diameter to a diameter
of 3 mm, wherein a reference vessel diameter is 3 mm and a diameter
stenosis of the diseased section of the blood vessel
pre-implantation is greater than 50%, wherein the vessel wall
comprises a plaque region including a necrotic core component, and
wherein as the stent degrades the area of the necrotic core
component as a percentage of a total area of the plaque region
decreases.
7. The method of claim 6, wherein when the stent is completely
degraded away, the area of the necrotic core component as a
percentage of the total plaque area is less than the percentage
post-stenting.
8. The method of claim 6, wherein the scaffolding is made from
PLLA.
9. A method of treating a diseased section of a blood vessel,
comprising: implanting a bioabsorbable polymeric stent comprising a
scaffolding composed of a pattern of struts at a diseased section
of a blood vessel to form a stented segment comprising the stent
and a vessel wall at the diseased section, wherein implanting
comprises deploying the stent from a crimped diameter to a diameter
of 3 mm, wherein a reference vessel diameter is 3 mm and a diameter
stenosis of the diseased section of the blood vessel
pre-implantation is greater than 50%, wherein the vessel wall
comprises plaque including a necrotic core component and a fibrous
component, wherein between post-implantation and when the stent is
completely or substantially degraded away the necrotic core
component as a percentage of the plaque decreases and the fibrous
component as a percentage of the plaque increases.
10. The method of claim 9, wherein the scaffolding is made from
PLLA with a crystallinity between 25-50%.
11. The method of claim 9, wherein the stent is completely and
substantially degraded away in at least two years.
12. The method of claim 9, wherein when the stent is completely and
substantially degraded away, the stented section exhibits
vasomotion.
Description
[0001] This is a continuation of application Ser. No. 12/510,076
filed Jul. 27, 2009, which is a continuation-in-part of application
Ser. No. 12/364,321 filed on Feb. 2, 2009, and application Ser. No.
12/506,121 filed on Jul. 20, 2009, and each of which are
incorporated by reference herein.
BACKGROUND OF THE INVENTION
[0002] Field of the Invention
[0003] This invention relates to methods of treatment of coronary
artery disease with bioabsorbable polymeric medical devices, in
particular, stents.
[0004] Description of the State of the Art
[0005] This invention relates to radially expandable
endoprostheses, that are adapted to be implanted in a bodily lumen.
An "endoprosthesis" corresponds to an artificial device that is
placed inside the body. A "lumen" refers to a cavity of a tubular
organ such as a blood vessel. A stent is an example of such an
endoprosthesis. Stents are generally cylindrically shaped devices
that function to hold open and sometimes expand a segment of a
blood vessel or other anatomical lumen such as urinary tracts and
bile ducts. Stents are often used in the treatment of
atherosclerotic stenosis in blood vessels. "Stenosis" refers to a
narrowing or constriction of a bodily passage or orifice. In such
treatments, stents reinforce body vessels and prevent restenosis
following angioplasty in the vascular system. "Restenosis" refers
to the reoccurrence of stenosis in a blood vessel or heart valve
after it has been treated (as by balloon angioplasty, stenting, or
valvuloplasty) with apparent success.
[0006] Stents are typically composed of scaffolding that includes a
pattern or network of interconnecting structural elements or
struts, formed from wires, tubes, or sheets of material rolled into
a cylindrical shape. This scaffolding gets its name because it
physically holds open and, if desired, expands the wall of the
passageway. Typically, stents are capable of being compressed or
crimped onto a catheter so that they can be delivered to and
deployed at a treatment site.
[0007] Delivery includes inserting the stent through small lumens
using a catheter and transporting it to the treatment site.
Deployment includes expanding the stent to a larger diameter once
it is at the desired location. Mechanical intervention with stents
has reduced the rate of restenosis as compared to balloon
angioplasty. Yet, restenosis remains a significant problem. When
restenosis does occur in the stented segment, its treatment can be
challenging, as clinical options are more limited than for those
lesions that were treated solely with a balloon.
[0008] Stents are used not only for mechanical intervention but
also as vehicles for providing biological therapy. Biological
therapy uses medicated stents to locally administer a therapeutic
substance. The therapeutic substance can also mitigate an adverse
biological response to the presence of the stent. Effective
concentrations at the treated site require systemic drug
administration which often produces adverse or even toxic side
effects. Local delivery is a preferred treatment method because it
administers smaller total medication levels than systemic methods,
but concentrates the drug at a specific site. Local delivery thus
produces fewer side effects and achieves better results.
[0009] A medicated stent may be fabricated by coating the surface
of either a metallic or polymeric scaffolding with a polymeric
carrier that includes an active or bioactive agent or drug.
Polymeric scaffolding may also serve as a carrier of an active
agent or drug.
[0010] The stent must be able to satisfy a number of mechanical
requirements. The stent must be capable of withstanding the
structural loads, namely radial compressive forces, imposed on the
stent as it supports the walls of a vessel. Therefore, a stent must
possess adequate radial strength. Radial strength, which is the
ability of a stent to resist radial compressive forces, is due to
strength around a circumferential direction of the stent.
[0011] Once expanded, the stent must adequately provide lumen
support during a time required for treatment in spite of the
various forces that may come to bear on it, including the cyclic
loading induced by the beating heart. For example, a radially
directed force may tend to cause a stent to recoil inward. In
addition, the stent must possess sufficient flexibility to allow
for crimping, expansion, and cyclic loading.
[0012] The treatment of coronary artery disease with a stent may
require the presence of the stent only for a limited period of
time. During or part of this limited time a healing process takes
place which includes changes in the structure of the vessel wall,
referred to as remodeling. After the healing process is completed,
the presence of the stent is no longer necessary.
[0013] Coronary stents made from biostable or non-erodible
materials, such as metals, have become the standard of care for
percutaneous coronary intervention (PCI) since such stents have
been shown to be capable of preventing early and later recoil and
restenosis. However, a stent made out of such biostable material
retains is mechanical or structural integrity and remains at the
implant site indefinitely unless it is removed by intervention or
is dislodged. Intervention presents risks to the patient and
dislodgement can have significant adverse consequences on the
patient. Leaving the stent at the implant site permanently also has
disadvantages. One disadvantage is that the stented segment has the
compliance of the stent which is very different from that of
healthy vessel segment. Another drawback of such durably implanted
stents is that the permanent interaction between the stent and
surrounding tissue can pose a risk of endothelial dysfunction and
late thrombosis.
SUMMARY OF THE INVENTION
[0014] Various embodiments of the present invention include a
method of treating a diseased section of a blood vessel,
comprising: implanting a bioabsorbable polymeric stent comprising a
scaffolding composed of a pattern of struts at a diseased section
of a blood vessel to form a stented segment comprising the stent
and a vessel wall at the diseased section, wherein the vessel wall
comprises a plaque region including a necrotic core component and a
fibrous component, wherein at pre-implantation and
post-implantation of the stent the necrotic core component is in
contact with a blood-contacting surface of the vessel wall, and
wherein as the stent degrades the fibrous component becomes
positioned between the necrotic core component and the
blood-contacting surface of the vessel wall so that the necrotic
core component is not in contact with the blood-contacting
surface.
[0015] Further embodiments of the present invention include a
method of treating a diseased section of a blood vessel,
comprising: implanting a bioabsorbable polymeric stent comprising a
scaffolding composed of a pattern of struts at a diseased section
of a blood vessel to form a stented segment comprising the stent
and a vessel wall at the diseased section, wherein the vessel wall
comprises plaque region including a necrotic core component, and
wherein as the stent degrades the area of the necrotic core
component as a percentage of the total plaque region area
decreases.
[0016] Additional embodiments of the present invention include a
method of treating a diseased section of a blood vessel,
comprising: implanting a bioabsorbable polymeric stent comprising a
scaffolding composed of a pattern of struts at a diseased section
of a blood vessel to form a stented segment comprising the stent
and a vessel wall at the diseased section, wherein the vessel wall
comprises plaque, wherein the volume and area of the plaque
increases during a first time period after implantation and then
decreases during a second time period after the first time period,
wherein the scaffolding is completely or substantially absorbed by
the end of the second time period.
[0017] Other embodiments of the present invention include a method
of treating a diseased section of a blood vessel, comprising:
implanting a bioabsorbable polymeric stent comprising a scaffolding
composed of a pattern of struts at a diseased section of a blood
vessel to form a stented segment comprising the stent and a vessel
wall at the diseased section, wherein the vessel wall comprises
plaque including a necrotic core component and a fibrous component,
wherein between post-implantation and when the stent is completely
or substantially degraded away the necrotic core component as a
percentage of the plaque has decreased and the fibrous component as
a percentage of the plaque has increased.
[0018] Additional embodiments of the present invention include a
method of treating a diseased section of a blood vessel,
comprising: implanting a bioabsorbable polymeric stent comprising a
scaffolding composed of a pattern of struts at a diseased section
of a blood vessel to form a stented segment comprising the stent
and a vessel wall at the diseased section, wherein the vessel wall
comprises a region of plaque, and wherein a radius of curvature of
the surface of the plaque region during degradation of the stent is
higher than a radius of curvature of the surface of the plaque
region pre-implantation.
BRIEF DESCRIPTION OF THE DRAWINGS
[0019] FIG. 1A depicts a radial cross-section of an artery.
[0020] FIG. 1B depicts a section of an artery in three
dimensions.
[0021] FIG. 2 depicts a stent.
[0022] FIG. 3 depicts generic absorption curves for bulk-eroding
polymers.
[0023] FIG. 4A depicts a cross-section of a stent and delivery
system including a stent mounted on a balloon.
[0024] FIG. 4B shows the balloon of FIG. 4A and the stent in an
expanded configuration in apposition to walls of a blood
vessel.
[0025] FIG. 4C shows the stent of FIGS. 4A and 4B with the balloon
removed supporting a diseased section of the vessel at an expanded
diameter.
[0026] FIGS. 5A-C represent a schematic representation of an
exemplary embodiment depicting the properties of an implanted
bioabsorbable stent as a function of time.
[0027] FIGS. 6A-B depict a blow molding process of a polymer
tube.
[0028] FIG. 7 depicts a stent pattern.
[0029] FIGS. 8A-B depict images of a bioabsorable stent used in a
clinical study.
[0030] FIG. 9 is a flow chart that summarizes a population of a
clinical study.
[0031] FIG. 10 depicts gray-scale IVUS-VH images and the
corresponding radiofrequency processed images of a vessel of
patient before stenting, post-stenting, 6 months after stenting,
and 2 years after stenting.
[0032] FIG. 11 depicts IVUS images with tissue echogenicity that
are two-dimensional slices of an implant site of a single patient
at post-PCI and at 6 months follow-up.
[0033] FIG. 12 depicts IVUS three-dimensional images with tissue
echogenicity at post-PCI and at 6 months follow-up.
[0034] FIGS. 13A-B depicts serial assessment of stent struts by
OCT.
[0035] FIGS. 14A-D depict exemplary OCT images of sections of a
treated vessel at six months follow-up which have undergone
different-degrees of absorption.
[0036] FIGS. 15A-C depict OCT images of a treated vessel post-PCT,
at 6 months follow-up, and 2 years follow-up, respectively.
[0037] FIGS. 16A-B, 17A-B, and 18A-B are IVUS, IVUS-VH, and OCT
images, respectively, for one patient post-PCI and at 2 years
follow-up.
[0038] FIG. 19A depicts an OCT image of a section of a treated
vessel in which arrows indicate complete tissue coverage of a
strut.
[0039] FIG. 19B depicts an OCT image of a section of a treated
vessel in which arrows indicate incomplete tissue coverage of a
strut.
[0040] FIG. 20A shows angiography measurements for patients treated
with methergine and nitroglycerine at 2 year follow-up.
[0041] FIG. 20B shows angiography measurements for patients treated
with acetylcholine and nitroglycerin at 2 year follow-up.
[0042] FIG. 21 depicts the results of in vitro radial strength
testing of one arm of bioabsorbable stents manufactured for
clinical trial.
[0043] FIG. 22 is a schematic representation of the structure of
vulnerable plaque.
DETAILED DESCRIPTION OF THE INVENTION
[0044] Various embodiments of the present invention relate to a
bioabsorbable stent and methods of treatment of a blood vessel
afflicted with coronary artery disease with the bioabsorbable
stent. These embodiments include a stent and treatment with the
stent which is made from a bioabsorbable polymer or polymers. In
these embodiments, the stent is implanted at an afflicted site or
section in the vessel, interacts with the vessel in a manner
(described in detail below) that elicits time dependent healing
responses from the vessel, and eventually disappears or
substantially disappears from the section, which is healed.
Therefore, the stent and treatment results in healing of the
afflicted section without the associated disadvantages (as
described in detail below) of a biostable stent.
[0045] Coronary artery disease refers to a condition in which the
arteries that supply blood to heart muscle become hardened and
narrowed or stenotic. This is due to the buildup of cholesterol and
other material, called plaque, on their inner walls. Such narrowed
or stenotic portions are often referred to as lesions. Coronary
artery disease includes restenosis which refers to the reoccurrence
of stenosis. Although there are probably several mechanisms that
lead to restenosis of arteries, an important one is the
inflammatory response, which induces tissue proliferation around an
angioplasty site. The inflammatory response can be caused by the
balloon expansion used to open the vessel, or if a stent is placed,
by the foreign material of the stent itself.
[0046] A wall of a healthy blood vessel is essentially made up of
three distinct layers surrounding the lumen through which blood
flows, the outermost advantitia, the media, and the intima. FIG. 1A
depicts a radial cross-section of an artery and FIG. 1B depicts a
section of an artery in three dimensions showing the intima (A),
and the media (C), and the advantitia (E). The cells of the intima
are supported by the internal elastic membrane (B) that separates
the intima from the media. The external elastic membrane or lamina
(D) (EEM or EEL) is a concentration of elastic fibers at the inner
boundary of the adventitia and the media.
[0047] The intima layer is made up of a single layer of cells which
are fat in the middle and thin at the edges. In arteries, the
intima is an elastic membrane lining and includes a smooth
endothelium on its inner surface that is in contact with blood
flowing through the lumen. The media is the middle layer of the
walls of arteries and is composed of smooth muscle and elastic
fibers. The adventitia is the outermost layer of an artery. It is
primarily a muscular structure contained within fibers of collagen,
a strong protein which is also found in tendons and ligaments. The
adventitia is therefore a very important component responsible for
the inherent strength of the artery. A healthy section of a blood
vessel wall includes all of these layers. However, one or both of
the intima layer or endothelium in a diseased section of a blood
vessel can be damaged or may not be present.
[0048] A stent may include a pattern or network of interconnecting
structural elements or struts. FIG. 2 depicts a view of a stent
100. In some embodiments, a stent may include a body, backbone, or
scaffolding having a pattern or network of interconnecting
structural elements 105. Stent 100 may be formed from a tube (not
shown). FIG. 2 illustrates features that are typical to many stent
patterns including cylindrical rings 107 connected by linking
elements 110. The cylindrical rings are load bearing in that they
provide radially directed force to support the walls of a vessel.
The linking elements generally function to hold the cylindrical
rings together and do not contribute significantly to the support
of the lumen. The structural pattern in FIG. 2 is merely exemplary
to illustrate the basic structure of a stent pattern.
[0049] A stent such as stent 100 may be fabricated from a polymeric
tube or a sheet by rolling and bonding the sheet to form the tube.
A tube or sheet can be formed by extrusion or injection molding. A
stent pattern, such as the one pictured in FIG. 2, can be formed in
a tube or sheet with a technique such as laser cutting or chemical
etching. The stent can then be crimped on to a balloon or catheter
for delivery into a bodily lumen.
[0050] In general, a stent can be made partially or completely from
a biodegradable, bioabsorbable, or biostable polymer. A polymer for
use in fabricating a stent can be biostable, bioabsorbable,
biodegradable or bioerodable. Biostable refers to polymers that are
not biodegradable. The terms biodegradable, bioabsorbable, and
bioerodable are used interchangeably and refer to polymers that are
capable of being completely degraded and/or eroded when exposed to
bodily fluids such as blood and can be gradually resorbed,
absorbed, and/or eliminated by the body. The processes of breaking
down and absorption of the polymer can be caused by, for example,
hydrolysis and metabolic processes.
[0051] In general, in order to facilitate healing of a diseased
section of a vessel, the presence of a stent is necessary for only
for a limited period of time. Therefore, a stent made from a
biodegradable polymer is intended to remain in the body for a
duration of time until its intended function of facilitating
healing a diseased section of a blood vessel is completed. After
the process of degradation, erosion, absorption, and/or resorption
has been completed, no portion of the biodegradable stent, or a
biodegradable portion of the stent will remain at the treated
section of the blood vessel. In some embodiments, very negligible
traces or residue may be left behind.
[0052] Reducing degradation time to the minimum time required for
successful treatment is advantageous since it allows further
surgery or intervention, if necessary, on a treated vessel to occur
sooner. Additionally decreasing degradation time helps reduce the
risk of late thrombosis.
[0053] Chemical hydrolysis of the hydrolytically unstable backbone
in some polymers is the prevailing mechanism for the degradation of
a bioabsorbable polymer. Other mechanisms of degradation, such as
enzymatic attack and metabolic processes, can also contribute to
degradation. Polymer erosion can be ideally divided into "bulk
erosion" and "surface erosion." For ideal bulk erosion, polymer is
chemical degraded and material is lost from the entire polymer
volume.
[0054] Although a bulk eroding polymer degrades throughout its
volume, a device made from a bulk eroding polymer can still
maintain its mechanical properties (e.g., strength) and mechanical
or structural integrity while it degrades. FIG. 3 depicts generic
absorption curves for bulk-eroding polymers showing the sequence of
polymer molecular weight (C1), strength (C2), and mass reduction
(C3) during degradation. Journal of Craniofacial Surgery, (8)2:89,
1997. As illustrated in FIG. 3, the degradation of a bulk eroding
polymer generally occurs in two phases. In the first phase
illustrated by time period T1, water penetrates the bulk of the
device, preferentially attacking the chemical bonds in the
amorphous phase and converting long polymer chains into shorter
water-soluble fragments. The resulting decrease in molecular weight
is shown by C1 in FIG. 3. For a semi-crystalline polymer, there may
be a reduction in molecular weight with minimal loss in physical
properties, which is illustrated by time period Tp in FIG. 3. This
is because degradation occurs in amorphous phase initially and the
device matrix is still held together by the crystalline regions.
The reduction in molecular weight is soon followed by a reduction
in mechanical properties (C2), and then erosion or mass loss (C3).
The mass loss eventually results in loss of structural integrity
demonstrated by fragmentation of the device. In the second phase
illustrated by T2, enzymatic attack and metabolization of the
fragments occur, resulting in a rapid loss of polymer mass.
[0055] Embodiments of the present invention include a bioabsorbable
stent and methods of treatment of coronary artery disease with the
stent. In such embodiments, a stent scaffolding can be formed from
bioabsorbable material such as a bioabsorbable polymer. In
particular, the stent can include a scaffolding made of
bioabsorbable polymer that is designed to provide support to a
vessel lumen once it is expanded. Exemplary biodegradable polymers
include poly(L-lactide) (PLLA), poly(D-lactide) (PDLA),
polyglycolide (PGA), and poly(L-lactide-co-glycolide) (PLGA). With
respect to PLGA, the stent scaffolding may be made from PLGA
between 0-100% L-lactide, in particular, 85/15 PLGA (85% L-lactide,
15% glycolide).
[0056] Additionally, the stent can further include a therapeutic
coating or layer above all or a portion of the scaffolding. The
coating can be composed of a bioabsorable polymer with one or more
therapeutic agents dispersed or dissolved in the polymer. The
therapeutic agents can include, but are not limited to,
antiproliferatives, and anti-inflammatories.
[0057] Various embodiments of a method of treating diseased blood
vessel include implanting or deploying the stent at a diseased
site, section, or segment of a blood vessel. The diseased section
can have a lesion which has caused stenosis or narrowing of the
blood vessel. Implantation can be performed by positioning the
stent at the diseased section and expanding the stent in apposition
to the vessel walls which increases the diameter of the section of
the vessel.
[0058] In the case of a balloon-expandable stent, the stent is
secured to a balloon at the end of a catheter prior to delivery and
deployment. Once the secured stent is positioned at the diseased
section of the blood vessel, the balloon is expanded to deploy the
stent. FIG. 4A depicts a cross-section of a stent and delivery
system including a stent mounted on a balloon 125 at a distal end
of a catheter 130. Individual stent struts 120 are shown secured on
balloon 125. The stent and delivery system are positioned within a
blood vessel 135 with walls 140 at a diseased section 145. FIG. 4B
shows balloon 125 and the stent in an expanded configuration in
apposition to walls 140 of blood vessel 135. FIG. 4C shows the
stent with the balloon removed supporting the diseased section of
the vessel at an expanded diameter.
[0059] After deployment, the stent maintains patency of the
diseased section for a limited period of time until chemical
degradation results in degradation of the radial strength to the
point that the stent can no longer support the walls of the section
of the vessel. Unlike a non-erodible stent, the mechanical
properties, structural integrity, and mass of the bioabsorbable
stent at the stented segment are time dependent since they change
during the healing process. The bioabsorbable stent provides
patency to the stented segment for a finite period of time, the
radial strength of the stent deteriorates, making the stent unable
to continue to provide patency to the vessel walls. The loss of
radial strength is followed by a gradual decline of mechanical
integrity, gradual loss of mass from the stent, and eventually
disappearance of the stent from the stented segment.
[0060] An essential feature of the stent of the present invention
is the time dependent nature of the mechanical properties,
mechanical integrity, and mass loss. The initial stent support due
to the radial strength followed by its loss, the gradual loss of
stent mechanical integrity, and gradual mass loss from the stent
results in elicits vessel responses that allow the stented section
to heal. The stent is designed so that the above stent behaviors
are timed with respect to biological responses of the vessel to
allow the healing processes to occur. The healed state is different
from that of a permanently stented segment provided by a
non-erodible stent.
[0061] During the time period that the stent provides support or
maintains patency of the lumen, the stent opposes the inward radial
force imposed by the lumen walls, including the cyclic loading
induced by the beating heart. The stent must maintain or sustain
such patency for a period of time in spite of the degradation or
erosion of the stent body. An exemplary desired degree of patency
is no less than 50% of the deployed diameter of the stent. Thus,
the stent should have sufficient strength, stiffness (modulus), and
creep resistance to keep recoil to an acceptable level during a
given period. Recoil refers to a decrease in diameter of a stent
from a deployed diameter. Therefore, an erodible stent structure
must have the appropriate combination of mechanical properties and
degradation or erosion properties to provide patency for a
particular period.
[0062] In the embodiments of the present invention, the stent has
design inputs that result in functional outputs that elicit healing
responses of the vessel. The design inputs include, generally,
mechanical, chemical, structural, microstructural properties, and
the processing parameters that result in such properties. As used
herein, functional outputs refer generally to two broad categories:
(1) the behavior of the stent once implanted and (2) vessel outputs
associate with healing. The stent outputs are associated with,
elicit, or facilitate the biological responses and vessel-stent
interactions.
[0063] The stent and method of treatment of the present invention
has design inputs that elicit and facilitate vessel outputs that
correspond to biological responses and vessel-stent interactions
associated with healing. Specifically, the stent has design inputs
that result in stent outputs that elicit or facilitate the healing.
The vessel outputs are observable as measurements from clinical
studies and demonstrate healing of the afflicted section. The
changes collectively correspond to healing of the vessel.
[0064] The implantation of a bioabsorbable stent with arbitrary
design inputs will inevitably elicit biological responses which
change the vessel. However, an arbitrary set of design inputs will
not necessarily result in vessel outputs or changes that result in
one or more aspects of healing of a vessel. For example, a stent
may result in formation of an endothelial layer, but not result in
positive remodeling that allows for increased blood flow.
Additionally, the functional outputs of an arbitrary set of design
inputs can be harmful and possibly lethal to the patient. For
instance, the stent may have catastrophic failure that induces a
thrombo-embolitic event. The stent of the present invention elicits
biological responses that result in selected aspects of a healed
vessel including changes in lumen dimensions that allow increased
blood flow, endothelialization, complete or partial restoration of
natural vessel compliance and vasomotion. The biological responses
are elicited through a synergistic combination of a set of design
inputs of the stent.
[0065] Design inputs may be classified into several broad
categories including, but not limited to, stent scaffolding design,
material properties--stent scaffolding chemistry, material
properties--thermo-mechanical, material processing parameters, and
therapeutic coating properties. Table 1 provides a summary
exemplary design inputs for each category. The design inputs are
not limited to those listed in Table 1.
TABLE-US-00001 TABLE 1 Design Inputs. Design Input Category Design
Inputs Stent scaffolding geometry of the stent pattern design strut
width and thickness. Materials class of degradable polymer (e.g.,
aliphatic property- polyester, alpha hydroxy polyester) scaffolding
type of polymer (e.g., poly(L-lactide) chemistry inputs intrinsic
hydrolysis rate. Material molecular weight (MW) or molecular weight
properties- distribution of the scaffolding polymer thermo-
mechanical properties of scaffold: radial strength, mechanical
fracture toughness inputs Tg and Tm of scaffolding polymer Degree
of uniaxial (circumferential) or biaxial (circumferential and
longitudinal orientation) obtained from radial expansion and/or
axial elongation the degree or percent crystallinity of the
scaffolding polymer size and distribution crystallites in
scaffolding Material extrusion parameters processing blow-molding
parameters parameters e-beam sterilization parameters Therapeutic
the type of polymer coating design type of drug dose of drug in the
coating thickness of coating.
[0066] Table 2 list stent outputs and exemplary design inputs which
provide these outputs. The list of design inputs for each output is
not limited to those listed in Table 2. The geometry of the stent
pattern refers the arrangement of struts in the stent pattern (see
e.g., FIG. 7). The intrinsic hydrolysis rate refers to the rate of
chemical degradation of a particular type of polymer which is
governed by the chemical composition of the polymer. The molecular
weight distribution refers to weight average (Mw) or number average
molecular weight (Mn). Degree of uniaxial (circumferential) or
biaxial (circumferential and longitudinal orientation) is provided
by radial expansion or radial expansion/axial elongation,
respectively, of the polymer tube from which the stent is
fabricated. A measure of the degree of orientation is provided by
the degree or percent of radial expansion and axial elongation,
respectively. The extrusion parameters are the parameters of the
extrusion process for fabricating the polymer tube from which the
stent is fabricated. The parameters include the temperature in the
extruder, pull rate, and draw-down ratio and quenching temperature
as the extrudate exits the die at the end of the extruder. The blow
molding parameters include the temperature expansion and percent
radial expansion and percent axial elongations.
TABLE-US-00002 TABLE 2 Stent outputs-properties refer to
scaffolding, except for drug release profile. Stent Output Design
Inputs Mechanical class and type of polymer properties profile: Tg
radial strength uniaxial/biaxial orientation and degree of thereof
profile, fracture degree of crystallinity toughness size and
distribution of crystallites blow molding parameters molecular
weight e-beam parameters intrinsic hydrolysis rate Drug release
profile type of coating polymer type of drug dose of drug in the
coating thickness molecular weight Mechanical integrity class and
type of scaffolding profile uniaxial/biaxial orientation and degree
of thereof degree of crystallinity size and distribution of
crystallites blow molding parameters molecular weight e-beam
parameters intrinsic hydrolysis rate Erosion profile class and type
of scaffolding intrinsic hydrolysis rate size and distribution of
crystallites
[0067] The time dependent radial strength profile of the stent
includes an initial period after intervention in which the stent
maintains its radial strength to prevent negative remodeling of the
vessel which is then followed by a loss of radial strength. As
discussed in more detail below, the length of time the stent can
provide support is particularly important in the healing,
specifically, remodeling to provide increased vessel
dimensions.
[0068] The radial strength loss arises from degradation of strength
primarily in the bending regions of the stent scaffolding. These
regions degrade and eventually are unable to oppose the force
imposed by the vessel wall. The stent scaffolding then exhibits a
controlled recoil inward.
[0069] The radial strength of the stent during the support period
and the time dependent behavior (i.e., the loss of radial strength)
are collectively due to several design inputs. These design inputs
synergistically provide stent behavior that allows initial
remodeling of the vessel wall with support, followed by transfer of
the load to the vessel wall with further remodeling. The design
inputs are listed in Table 1 for stent scaffolding design,
materials property-scaffolding chemistry inputs, material
properties-thermo-mechanical inputs, and material processing
parameters.
[0070] Design inputs include the class and type of scaffolding
polymer, Tg, uniaxial/biaxial orientation and degree of thereof,
degree of crystallinity, size and distribution of crystallites,
blow molding parameters, molecular weight, e-beam parameters, and
intrinsic hydrolysis rate. The stent of the present invention can
be designed provide mechanical support between 1 and 4 months after
intervention. In the clinical studies of the BVS stent discussed
below, OCT images (FIG. 11) at 6 months indicate partial absorption
of stent struts which indicates that support by the stent ended
before 6 months.
[0071] The scaffolding polymer can be an aliphatic,
semicrystalline, degradable polymer with a Tg above human body
temperature, about 37.degree. C. The semicrystalline polymer is
processed to have a degree and nature of the crystallinity of a
semicrystalline polymer that provides a high strength and fracture
toughness. Although crystallinity provides strength to a polymer,
if the crystallinity is too high, stent scaffolding can become
brittle and be susceptible to fracture and failure once
implanted.
[0072] Additionally, to increase the strength to weight ratio of
the stent so that thinner struts can be used to reduce the stent
profile, a precursor to the stent, a polymer tube, is radially
expanded prior to forming a stent pattern in the tube. The polymer
tube is expanded using blow molding in a temperature range that is
expected yield a high nucleation density with small, well dispersed
crystallites. The degree of crystallinity, nature of crystallinity,
and radial orientation also increase the fracture toughness. These
characteristics combined, reduces or eliminates fracture and
failure of stent struts as the stent loses radial strength and
reduces the chance of catastrophic failure of the stent during
radial strength decline. The high fracture toughness allows for a
controlled loss of radial strength during support and during loss
of radial strength.
[0073] The molecular weight of the final product is designed to be
high enough to provide the required radial strength and also to
provide the time dependent behavior. Process parameters of process
steps that degrade the molecular weight are adjusted to reduce the
molecular weight degradation. These include the extrusion
temperature and radiation does of the electron beam
sterilization.
[0074] The drug release profile includes a release of
antiproliferative drug during smooth muscle proliferation (SMP),
which declines to zero to allow healing processes to occur.
Specific aspects include the amount drug release and the time
dependent drug release profile from the coating.
[0075] The stent of the present invention is designed to provide a
release profile which controls proliferation during smooth muscle
cell proliferation, but terminates soon enough to allow complete or
almost complete endothelialization prior to substantial mass loss
and mechanical integrity loss. "Almost complete" can correspond to
at least 90% of struts covered by an endothelial layer.
Specifically, the stent is designed to have a drug release profile
that declines to zero between 3-4 months after intervention. As
indicated below, stent is designed to allow for complete or almost
complete endothelialization of stent struts between 4 and 6 months
after intervention.
[0076] The design inputs include, but are not limited to, the type
of polymer, type of drug, dose of drug in the coating, thickness,
and molecular weight. The type of polymer is faster eroding than
the scaffolding polymer, and a high fracture toughness. An
exemplary polymer is an amorphous polymer (contributes to faster
erosion and higher fracture toughness) with a faster intrinsic
hydrolysis (for faster erosion), such as poly(DL-lactic acid). The
drug-polymer ratio is relatively high to reduce the required
thickness of the coating, for example, between 40-60 wt % drug. The
thickness of the coating is less than 3 microns and preferably less
than 2 microns. The low thickness also reduces the profile of the
stent which limits the likelihood of thrombosis. A molecular weight
lower than the scaffolding polymer contributes to faster erosion
rate. A weight average molecular weight of the coating is between
40,000 and 80,000.
[0077] Mechanical integrity refers to the size, shape, and
connectivity of the structural elements of the stent. For example,
the shape refers to the generally tubular shape of the stent formed
by the cylindrically-shape rings connected by the linking elements
of the pattern.
[0078] An initial loss of mechanical integrity occurs when some or
all of the linking elements have failed resulting in partial or
complete loss of connectivity between cylindrical rings. The
cylindrical rings can remain intact for a period of time and
maintain a circular shape. Further loss of mechanical integrity
occurs when there is a loss of connectivity between structural
elements in the cylindrical rings.
[0079] The stent of the present invention is designed to exhibit
features of mechanical integrity profile that are critical to
vessel outputs associated with healing. The mechanical integrity is
maintained a sufficient time to enable endothelialization (e.g.,
between 90-100% of struts covered) within 4-6 months after
intervention. Additionally, the stent is designed to provide a
gradual loss mechanical integrity after loss of radial strength to
provide a restoration of vessel compliance and vasomotion.
[0080] The mechanical integrity profile of the stent is due to
several design inputs which synergistically provide the profiles
that provide the vessel outputs associated with healing. These
mechanical integrity arises from several design inputs listed in
Table 1. These stent inputs relate to stent scaffolding design,
materials property-scaffolding chemistry inputs, material
properties-thermo-mechanical inputs, and material processing
parameters.
[0081] Stent inputs include class and type of scaffolding polymer,
uniaxial/biaxial orientation and degree thereof, degree of
crystallinity, size and distribution of crystallites, blow molding
parameters, molecular weight, e-beam parameters, and intrinsic
hydrolysis rate. The chemistry and molecular weight of the
scaffolding polymer contribute to the timing of the mechanical
integrity and erosion profile.
[0082] As illustrated in FIG. 2, mass loss can be insignificant
even after strength is lost and after a significant loss of
molecular weight. The mechanical integrity and erosion profiles
depend at least on the intrinsic hydrolysis rate of the polymer
(class and type) and on the initial molecular weight of the
scaffolding. The mechanical integrity profile also depends the
strength and fracture toughness and factors related to these such
as uniaxial/biaxial orientation and degree of thereof, degree of
crystallinity, size and distribution of crystallites, and blow
molding parameters. The strength and fracture toughness and the
contributing factors influence the manner and timing of the
breaking put of the stent pattern. The high strength and fracture
toughness enables the scaffolding to hold its shape and break apart
in a controlled manner with a low risk of thrombosis. For example,
as discussed in more detail below, the stent scaffolding should
maintain mechanical integrity until there is complete or almost
complete endothelialization.
[0083] Furthermore, it is expected that the size and distribution
of crystallites contribute to gradual loss of mechanical integrity.
The crystallites act as tie points or crosslinks that can help hold
together the polymer even after molecular weight loss.
[0084] Additionally, the integrity of the rings of the stent
pattern should be maintained until at least partially incorporated
into the vessel wall. The stent struts of the rings remain
connected until incorporated into a vessel wall by the endothelial
layer. Additionally, the formation of the endothelial layer and the
manner and timing of stent pattern break-up depends on the
cross-sectional size of the struts (i.e., the thickness and width
of the struts) and the fracture toughness and strength of specific
portions of the pattern. A larger strut cross-section provides
higher strength and delays pattern break-up, however, the struts
with a larger cross-section present a large profile to blood flow
and take longer for the endothelial layer to incorporate. With
regard to the latter, the links can be designed to fail first,
leaving the rings to fail after endothelialization and absorption
into vessel wall.
[0085] In the clinical studies of the BVS stent discussed below,
reduction in molecular weight and mass had occurred to such an
extent 2 years after intervention that struts were no longer
recognizable by intravascular ultrasound, leaving behind few
visible features. A third of stents were no longer discernible by
OCT.
[0086] The stent is designed to have an erosion profile such that
there is insignificant mass loss until after loss of the radial
strength, mechanical integrity, and endothelialization. The stent
is further designed so that there is complete absorption or
incorporation of remaining struts into the vessel walls by 2 years
after intervention. The erosion profile is provided by a
synergistic combination of several design inputs. The class and
type of polymer determine the rate at which the molecular weight,
strength, and mass is lost by an arbitrary volume of polymer in the
scaffolding. The erosion profile also depends on the crystallinity
and size and distribution of crystallites. This is because
hydrolysis is faster in amorphous regions than crystalline regions.
In the clinical studies of the BVS stent discussed below, the OCT
data in FIG. 15C suggest that the struts are almost or completely
dissolved at 2 years.
[0087] FIGS. 5A-C depict a schematic representation of exemplary
time dependent behavior of a bioabsorbable stent after intervention
at an afflicted section of a vessel. In addition, FIGS. 5A-C also
show expected biological responses of the vessel to the stent as a
function of time. Although the time scale shown is exemplary, the
time dependence of stent behavior is a qualitative representation
of the behavior the bioabsorbable stent of the present invention
which elicits healing of the afflicted section.
[0088] Each of FIGS. 5A-C shows the time dependence of the stent
properties, the radial strength, drug release, mechanical
integrity, and erosion or mass loss. The radial strength of the
stent is maintained for a period of time (in this case, between
2.5-3 months after intervention) after intervention during which
the stent supports the vessel walls. The stent then experiences a
rapid deterioration in radial strength, due to molecular weight
loss, and can no longer support the lumen walls (in this case,
about 3 months after intervention). The drug release is maintained
at a relatively constant level after intervention (in this case,
between 1-1.5 months after intervention) followed by a relatively
rapid decline to zero (in this case, between 3-4 months after
intervention). The mechanical integrity is maintained at a
relatively constant level for a period of time after intervention
(in this case, about 3-4 months after intervention) followed by a
gradual decline until a complete loss at a time greater than 6
months. The period of mechanical integrity retention is longer than
radial strength retention and the rate of decline of mechanical
integrity is more gradual.
[0089] FIGS. 5A-C shows that that is insignificant mass loss until
loss of the radial strength and mechanical integrity. Complete mass
loss or full bioabsorption occurs after about two years after
intervention.
[0090] There are several phases of biological response and vessel
changes due to the intervention of the stent. The time period from
intervention to about 1-3 months after intervention is referred to
as the acute phase. FIG. 5A depicts two biological responses to the
stent that occur during this phase, platelet deposition and
leukocyte or white cell recruitment. These biological responses can
dissipate quickly if there is growth of cellular layers over the
stent. In FIG. 5A, platelet deposition peaks after two weeks and
decays to a negligible level at about two months. Leukocyte
recruitment peaks at slightly after one month, decreases rapidly,
and trails off to zero at about five months.
[0091] FIG. 5B depicts additional biological responses during the
acute phase, smooth muscle cell proliferation (SMP) and matrix
deposition. SMP occurs at the inner surface of the vessel wall in
the stented section. The exemplary profile in FIG. 5B shows that
the smooth muscle cell proliferation reaches a peak between one and
two months and then decreases to negligible levels at about five
months. Smooth muscle cell proliferation can be explained with
reference to the structure of an arterial wall. During smooth
muscle cell proliferation, smooth muscle cells migrate from the
media layer to the vessel wall surface and proliferate to form a
neointima layer. Smooth muscle cell proliferation is expected to
occur during a time period up to about three months after
implantation of the stent. As explained below, smooth muscle cell
proliferation must be controlled since it can lead to excessive and
undesirable narrowing of the stented segment.
[0092] Matrix deposition involves deposition of collagen and
elastin in the neointima layer, reinforcing the layer which enables
it to provide mechanical support. Matrix deposition is a key
component of the remodeling process. Remodeling refers to a
biological response that results in modification of the neointima
layer formed from smooth muscle cell proliferation. The
modification facilitates a restoration of normal function of the
vessel. The remodeled neointima layer includes a smooth muscle cell
matrix banded together with elastin and collagen. The remodeling
process is expected to start within about a week after intervention
and can occur up to about 6 months or beyond 6 months after
implantation.
[0093] Another biological response to the stent includes formation
of an endothelial layer over the neointima layer and the stent.
FIG. 5C, which depicts cumulative endothelialization as a function
of time, shows that endothelialization starts shortly after
implantation and reaches a maximum just before three months. A
stent that allows for complete or almost complete endothelial
coverage in an appropriate time frame is essential since the
endothelial layer facilitates healing of the diseased section.
[0094] The stent of the present invention has design inputs and
stent outputs that elicit vessel outputs that are associated with
healing. The specific features of these vessel outputs are
discussed below. Table 3 lists vessel outputs and the stent outputs
and design inputs that correspond to the vessel outputs.
TABLE-US-00003 TABLE 3 Vessel Outputs. Vessel Output Stent outputs
Remodeling with support to Radial strength profile prevent elastic
recoil Controlled Smooth Muscle Drug release profile Cell
proliferation Endothelialization Drug release profile Mechanical
integrity Erosion profile Restoration of vessel compliance Radial
strength profile and homogeneity Mechanical integrity profile Time
dependent changes in vessel Drug release profile
dimensions-diameter, area, volume, etc. Mechanical integrity
Erosion profile Restoration of vasomotion Radial strength profile
Mechanical integrity profile Erosion profile
[0095] As indicated above, an essential feature of the stent is the
time dependent nature of the mechanical properties, specifically,
the stent provides support through its radial strength for an
initial period after intervention. This initial period of support
is necessary to allow sufficient time for the neointimal layer to
remodel at an increased diameter. Without sufficient remodeling
time with support, the vessel will be unable to heal properly,
i.e., to be restored to a natural functioning state. However, it is
also essential for the radial strength decline in a controlled
manner at some point so that the vessel can complete the healing
process and revert to the natural functioning state.
[0096] The stent of the present invention is designed to provide
support or patency to the stented section of a vessel for a limited
period of time. The support of the stent prevents elastic recoil
and negative remodeling, referring to remodeling at a decreased
diameter. The stent provides a scaffold to maintain a circular
lumen while the vessel remodels and molds itself to the stented
diameter. Negative remodeling can result in vessel dimensions, such
as diameter and lumen area, after stent absorption that is
substantially less than a normal, healthy vessel.
[0097] The support is primarily due to the radial strength of the
stent so the stent is designed to have a radial strength that is
sufficient to provide this support. The stent is additionally
designed to lose radial strength after the period of support so
that it is no longer able to support the lumen. The loss of radial
strength represents a transition of the load bearing from the stent
to the partially remodeled neointimal layer, allowing completion of
the healing process. Additionally, the stent is designed to
maintain radial strength and lose radial strength without
catastrophic failure which could result in an adverse thrombotic
event.
[0098] After a period of time of mechanical support, the radial
strength of the stent rapidly deteriorates to a degree that the
stent can no longer provide support to the vessel walls. In some
embodiments, the time period of support can be less than 1, less 3
months, or less than 6 months after intervention. However, the
remodeling that occurs during the period of mechanical support is
sufficient to allow the vessel to heal. In particular, the vessel
walls can maintain an increased diameter after the stent disappears
from the vessel.
[0099] Although smooth muscle cell proliferation is an essential
feature of the remodeling process, it is necessary to control the
proliferation of smooth muscle cells. In the absence of control,
the smooth muscle cell layer can be undesirably thick, causing
restenosis. Therefore, the stent is designed to release of an
antiproliferative agent from a therapeutic coating layer over the
stent scaffolding to control the smooth muscle cell proliferation.
The therapeutic agent release can occur up to two or four months
from intervention.
[0100] The time period of release is critical. It is important to
control the proliferation of the smooth muscle cells with the
antiproliferative agent. On the other hand, it is also important,
as explained below, to achieve partial or complete
endothelialization of the stent surface early in the healing
process. In particular, it is essential to achieve partial or
complete endothelialization prior to substantial loss of mechanical
integrity and mass loss since the antiproliferative agent tends to
inhibit endothelialization.
[0101] The endothelialization plays a critical role in the healing
process with a bioabsorbable stent. Both the degree of
endothelialization and timing of the endothelialization with
respect the stent behavior are crucial outputs. Endothelialization
refers to coverage of a surface with a layer of endothelial cells.
Complete or almost complete endothelialization of the vessel wall
and stent struts is essential to prevent thrombosis associated with
blood contacting stent surfaces, incomplete strut apposition
(persistent or late-acquired), and dislodgement of stent material.
Additionally, the timing of the endothelialization with respect to
mechanical integrity loss and mass loss is also an important aspect
of the healing process.
[0102] The presence of a blood-contacting surface of a foreign body
regardless of level of hemo-compatibility, such as a stent,
presents the risk of thrombosis. In general, an endothelial layer
plays a crucial role in reducing or preventing vascular thrombosis
and intimal thickening. Specifically, the endothelial layer reduces
or prevents deposition of proteins on the vessel wall or stent
struts. Such deposition can contribute to or increase risk of
thrombosis. Therefore, early and complete endothelialization of the
vessel wall and stent are essential.
[0103] Incomplete stent apposition can creates a risk of
thrombosis. Incomplete apposition can occur at intervention and
persist for several months. Additionally, stent struts can dislodge
from complete apposition with the vessel wall after intervention
and is referred to as late-acquired stent apposition (LAISA). In
either case, stent struts protrude into the lumen, presenting a
obstacle to blood flow and risk of a thrombo-embolitic event.
Incomplete apposition can occur in nonerodible stents and before
loss of radial strength and mechanical integrity in a bioabsorbable
stent. However, LAISA can become pervasive with a bioabsorbable
stent when radial strength and mechanical integrity decline. With
the decline of radial strength, the vessel can push back on the
stent struts, resulting in creep of the struts inward. The
development of an endothelial layer over the stent and vessel wall
reduces or prevents such adverse events since the biocompatible
endothelial layer covers dislodged struts.
[0104] In the BVS clinical studies, OCT images in FIGS. 13A-B show
that persistent incomplete and late acquired incomplete apposition
detected at 6 months and previously reported, were no longer
detectable at 2-year follow up. At 2 years, there was a smooth
appearance of the endolurninal lining without strut malapposition
since struts have been absorbed. Since no thrombo-embolitic events
were detected, it is believed that the endothelial lining inhibited
such events.
[0105] Additionally, as mechanical integrity declines and mass is
lost from the stent, there is increased risk of stent material
dislodging and completely separating from the vessel wall into the
blood stream, which can cause a thromo-embolitic event. Endothelial
layer coverage reduces or prevents such complete separation as
mechanical integrity and mass is lost. Therefore, it is crucial for
the complete or almost endothelialization prior to substantial loss
of mechanical integrity and mass loss. Such substantial loss of
mechanical integrity can include pieces of the stent struts.
[0106] As indicated above, the timing of the complete or almost
complete endothelialization relative to radial strength loss,
mechanical integrity loss, and mass loss is crucial. Complete or
almost complete endothelialization should occur between 4 and 6
months to reduce the risk of or avoid the thrombo-embolitic events
associated with LAISA and dislodgement of material in the vessel.
Endothelialization prior to substantial mass is important since
release of acidic hydrolytic degradation products inhibit
endothelialization.
[0107] The stent of the present invention is designed to have drug
release, mechanical integrity, and erosion profiles that enable
endothelialization which reduces or prevents thrombo-embolitic
events, in particular, those associated with LAISA and material
dislodgement. The antiproliferative drug release rapidly declines
to zero by 3 to 4 months after implantation so as not to interfere
with endothelial growth. Additionally, the mechanical integrity
remains substantially intact until about _2.sub.-- to _6.sub.--
months until the complete or almost complete endothelialization.
The erosion profile is such that significant mass loss starts only
after endothelialization is complete or almost complete.
[0108] The growth of the endothelial layer is facilitated by the
biocompatibility of stent material. Materials that have no or a low
degree of cytotoxicity are biocompatible and can result in rapid
endothelial growth and healing. Biocompatible polymers include, but
are not limited to, poly(L-lactide), poly(DL-lactide),
polyglycolide, poly(L-lactide-co-glycolide) and polycaprolactone.
In a human patient, endothelial layer growth can occur between
post-stenting to 3 months, or up to six months, or more than six
months after implantation. In some embodiments, at least 90%, 95%,
or at least 99% of stent struts can be covered by an endothelial
layer by six months after implantation.
[0109] As indicated in Table 3, the stent of the present invention
allows for restoration of vessel compliance and homogeneity to that
the vessel. The compliance of a segment of a vessel is the change
in luminal area per unit change in distending pressure in the
vessel. A segment refers to a longitudinal section of a vessel with
or without the presence of a stent. Thus, as used herein,
compliance of a stented segment is the compliance of the composite
structure that includes both the stent and the vessel. In the
absence of a stent, the segment has the compliance of the vessel
walls.
[0110] OCT imaging data from the clinical studies of the BVS stent
show absorption of the stent into artery walls and that the blood
vessel lining of arteries treated with the stent looks more uniform
after two years than it did immediately post-treatment.
[0111] The compliance of a stented segment in a treatment with the
stent of the present invention changes with time. The change in the
compliance is due both to the time dependence of stent properties
and to the changes in the vessel wall with time. As discussed
above, the radial strength, mechanical integrity, and mass of stent
change as a function of time. The changes in the vessel with time
are primarily due to the remodeling of the vessel wall.
[0112] The compliance at intervention is very low due to the
strength and stiffness of the stent. The compliance eventually
converges to that of the natural compliance of the vessel when it
is healed. The compliance of the stented segment is dominated by
the stent during the period of support by the stent. This is
followed by a rapid increase in compliance when the radial strength
of the stent declines. As the mechanical integrity of the stent
declines and the stent gradually erodes, the compliance of the
stented segment gradually is restored to that of a healed
vessel.
[0113] As indicated above, the vessel wall is undergoing remodeling
during the period of support during which the stent dominates the
compliance. However, the vessel walls are also undergoing
remodeling as the compliance of the stented segment changes and
converging to that of the healed vessel. Even after the decline of
radial strength and before the loss of structural integrity, the
scaffolding can still restrict or inhibits freedom of movement of
the vessel wall in response to change of pressure in the vessel.
The degree of restricting gradually decreases as the scaffolding
breaks up and erodes. Since the compliance is converging to the
vessel wall, the vessel wall is undergoing movement or vasomotion
as it is remodeling.
[0114] The clinical palpography data for the BVS stent presented
below in Table 10 demonstrate that compliance convergence. The
deformability of the stented segment increased significantly from
intervention to 6 months follow-up and then increased slightly from
6 months to 2 years follow-up.
[0115] The stent of the present invention further provides for
partial or complete restoration of vasomotion in a healed state.
Vasomotion refers to rhythmic oscillations in vascular tone caused
by local changes in smooth muscle. The stent outputs and design
inputs that provide convergence of the compliance of the stented
segment further allow for restoration of vasomotion.
[0116] Based on clinical results presented below for the BVS stent
in FIGS. 20A-B, previously stented portion of arteries demonstrated
the ability to expand and contract in a manner similar to a vessel
that has never been stented. Additionally, the OCT data of the BVS
clinical data showed an optically homogeneous vessel wall structure
that, taken together with the documented restoration of vasomotion,
suggests healing of the artery.
[0117] The gradual convergence of the compliance is facilitated by
the mechanism of mechanical integrity loss. As discussed above, the
circular rings of the pattern can be decouple and remain intact as
the links fail. The stented segment returns to a natural state as
the circular rings gradually break apart, incorporated into the
vessel wall, and are slowly absorbed. IVUS with echogenicity, IVUS
with virtual histology, and OCT results for the clinical study of
the BVS stent indicate that by 2 years, the BVS stent was
incorporated into the vessel wall and bioabsorbed.
[0118] The convergence of the compliance of the stented segment to
that of a vessel is collectively due to several design inputs. The
stent inputs, listed in Table 1, include stent scaffolding design,
materials property-scaffolding chemistry inputs, material
properties-thermo-mechanical inputs, and material processing
parameters.
[0119] Design inputs such as the chemistry and initial molecular
weight of the scaffolding provide the degradation profile of the
molecular weight, radial strength, and mass. The stent is designed
to have high radial strength and fracture toughness so that the
stent can lose radial strength without catastrophic failure and so
that the rings can remain intact, be absorbed in the vessel wall,
and gradually disintegrate. The design inputs that contribute to
high radial strength and fracture toughness are discussed above.
Additionally, the stent scaffolding is designed to have hoop or
circumferential strength that is greater than the strength
transverse to the circumferential direction. This is due to greater
preferential polymer chain orientation in the circumferential
direction than the transverse direction. The greater preferential
orientation produced through. This difference in strength
facilitates failure of links that decouple adjacent rings. This
decoupling allows movement of the vessel wall similar to an
unstented vessel as the vessel heals.
[0120] As indicated in Table 3, the time dependent stent changes in
vessel dimensions of a vessel are related to stent outputs
including drug release profile, radial strength profile, mechanical
integrity profile, and erosion profile. The drug release profile
controls smooth muscle cell proliferation and prevents restenosis.
The drug release decays early enough to allow endothelialization.
The radial strength profile provides support for a period of time
to prevent negative remodeling of the vessel wall followed by a
rapid loss of radial strength. The mechanical integrity declines
gradually, during which further remodeling takes place. Thus, each
of these stent outputs influences the vessel dimension as a
function of time.
[0121] Based on the clinical results presented below for the BVS
stent, the radial strength was lost between intervention and 6
months. There was a significant loss of mechanical integrity by 6
months follow-up. The IVUS data in Table 8 show that the minimal
luminal area, average luminal area, and lumen volume decrease from
intervention and 6 months. The IVUS data showed these parameters
increased between 6 months and 2 years after intervention.
[0122] OCT data in Table 11 indicates a decrease in mean lumen
area, minimal lumen area, lumen volume, and mean lumen diameter
from intervention to 6 months follow-up. Like the IVUS results,
between 6 months and 2 years, OCT showed an increase in these
quantities.
[0123] The design inputs that provide the above-described stent and
vessel outputs are as described below. The Tg of the scaffolding
polymer is preferably between 10 and 30.degree. C. above Tg to
insure stiffness at human body temperature. For example, PLLA has a
Tg of about 60.degree. C. The degree of crystallinity that
contributes to radial strength is 25-50%. The range of radial
expansion is 300 to 500%, as defined below. The temperature range
of radial expansion depends on the particular polymer. However, the
temperature range is in a range that is less than
Tg+0.6.times.(Tm-Tg). Additionally, the polymer tube from which the
stent is made is extruded so that it has a crystallinity of less
than 15% so most of the crystallinity of the stent scaffolding can
be generated in the manner described above. The weight average
molecular weight range of the scaffolding struts is between 50,000
and 300,000. The range of width and thickness of the struts is 100
to 200 microns. The range of cross-sectional area is about 17,000
to 40,000 square microns.
[0124] As indicated above, hardening and narrowing of arteries is
attributed to the buildup of cholesterol and plaque on the inner
walls of a vessel. The morphology and composition of plaque can be
characterized as having four tissue components: necrotic core,
dense calcium, fibro-fatty, and fibrous tissue. The volume and
cross-sectional area of these various components can be measured by
analytical techniques that are described below. The plaque region
can extend completely or partially around the circumference of the
vessel wall.
[0125] The necrotic core of the plaque is formed inside of the
vessel wall when it starts to undergo apoptosis or necrosis. The
necrotic core has essentially no living cells and mainly includes
cholesterol debris, such as cholesterol crystals and dead cell
debris. The necrotic core is known to be vulnerable to rupture
which can result in blood clot formation.
[0126] Fibro-fatty plaque is characterized by an accumulation of
lipid-laden cells underneath the intima of the arteries, typically
without narrowing the lumen due to compensatory expansion of the
bounding muscular layer of the artery wall. Beneath the endothelium
there is a "fibrous cap" covering the atheromatous "core" of the
plaque.
[0127] Fibrous plaque is localized in the intima, within the wall
of the artery resulting in thickening and expansion of the wall
and, sometimes, spotty localized narrowing of the lumen. The
fibrous plaque contains collagen fibers (eosinophilic),
precipitates of calcium (hematoxylinophilic), and, rarely,
lipid-laden cells. The fibrous component forms a fibrous cap that
separates the other plaque components from the blood pool. The
fibrous cap can become thinner resulting in exposure of unstable
necrotic core close to or at the surface of the plaque.
[0128] The biodegradable stent of the present invention results in
changes in properties of plaque with time as the stent degrades.
This has been demonstrated by clinical results presented herein.
Stent outputs and design inputs provide for the time dependent
behavior of these properties.
[0129] The time dependent behavior includes the change in the
luminal area and volume of the plaque as the stent degrades between
implantation and 2 years, at which time the stent is almost
completely eroded away. The plaque volume and area have been shown
to increase over the first 6 months after implantation and then
decrease between 6 months and 2 years, when the scaffolding is
completely or substantially absorbed.
[0130] The time-dependent behavior also includes the change in the
plaque geometric morphology parameters and the composition with
time. These changes are demonstrated by the clinical results
presented herein. The time-dependence of plaque morphology
parameters and composition have been demonstrated by clinical
results presented herein, for example, as shown in FIG. 10.
[0131] The parameters include the ratio of the thickness of the
fibrous cap to the necrotic core. The clinical results have shown a
regression of necrotic core as the stent degrades. Specifically, a
vessel wall can have a necrotic core component at pre- and
post-implantation that is in contact with a blood-contacting
surface of the vessel wall. As the stent degrades the necrotic core
regresses from the surface and the fibrous component becomes
positioned between the necrotic core component and the
blood-contacting surface of the vessel wall. As a result, the
unstable necrotic core component is no longer in contact with the
blood-contacting surface.
[0132] Another time-dependent plaque morphology parameter is the
area of the necrotic core component as a percentage of the total
plaque area. The clinical results show that as the stent degrades
the area of the necrotic core component as a percentage of the
total plaque region area decreases.
[0133] Another time-dependent plaque morphology parameter is
necrotic core component as a percentage of the plaque. The clinical
results show that between implantation and when the stent is
completely or substantially degraded away, the necrotic core
component as a percentage of the plaque has decreased and the
fibrous component as a percentage of the plaque has increased.
[0134] Another plaque morphological parameter that changes with
time is the shape of the plaque region. The implantation of the
stent results in a change in the shape of a plaque region which is
more stable and less susceptible to rupture. When the stent is
implanted, the plaque region is redistributed, resulting in a
change in shape. In addition, the redistribution changes the shape
of the plaque region. The pressure exerted by the stent on the
plaque region and vessel wall changes the radius of curvature of
the plaque region. In particular, the surface of the plaque region
is flattened. In other words, the radius of curvature of the plaque
region is increased by the deployment of the stent. The pre- and
post-stenting IVUS-VH images shown in FIG. 10 demonstrate this
change in shape. This flattened profile of the plaque region is
more stable and less vulnerable to rupture since there is lower
stress at the corners of the plaque. This lower stress makes the
plaque less susceptible to rupture.
[0135] Additionally, the plaque properties change with time as the
stent degrades which allows the plaque to maintain the more stable
shape even after the stent degrades. Specifically, it is believed
that the vessel wall, including the plaque region, undergoes
remodeling as the stent degrades. Due to the remodeling, the
flattened shape or increased radius of curvature is maintained,
even after the radial strength of the stent declines and is no
longer able to support the vessel walls. Furthermore, the flattened
shape or increased radius of curvature is maintained after the
stent loses mechanical integrity and even after the stent erodes
away completely from the section of the vessel. The IVUS-VH results
shown in FIG. 10 show that the flattened shape is maintained at 2
years, at which the stent is almost completely eroded away.
[0136] Stent outputs and design inputs result in the time dependent
behavior of the luminal area and plaque morphology parameters. The
time-dependent behavior of the luminal area arises from stent
outputs including the mechanical integrity profile and the radial
strength profile. The time-dependent plaque geometric morphology
parameters are a result of the mechanical integrity profile, radial
strength profile, stent design, erosion profile. The remodeling of
the plaque region is provided by the stent outputs and design
inputs as described for remodeling of the vessel wall.
[0137] The bioabsorbable stent of the present invention may be used
to treat a section of a vessel having vulnerable plaque. A
vulnerable plaque is an unstable collection of white blood cells
(primarily macrophages) and lipids (including cholesterol) in the
wall of an artery which is particularly prone to produce sudden
major problems, such as a heart attack or stroke. Vulnerable plaque
has a thin fibrous cap and a large and soft lipid pool underlying
the cap. FIG. 22 is a schematic representation of the structure of
vulnerable plaque. The upper panel shows the structure of unstable
or vulnerable plaque. The fibrous cap is thin contains few VSMCs
and collagen fibers. There are a large number of macrophages
particularly at the shoulder region where the cap meets the healthy
endothelium. There is a large lipid core. The lower panel shows a
stable plaque with a thick fibrous cap and a small lipid core.
There many VSMCs and collagen fibers. The characteristics of
vulnerable plaque together with the usual hemodynamic pulsating
expansion during systole and elastic recoil contraction during
diastole contribute to a high mechanical stress zone on the fibrous
cap of the atheroma, making it prone to rupture.
[0138] In such embodiments, the bioabsorbable stent may be deployed
at a section of the blood vessel that has a vulnerable plaque
region. The deployment of the stent would increase the radius of
curvature of the vulnerable plaque region. It is expected that as
the stent degrades the fibrous cap remodels and would be capable of
maintaining an increased radius of curvature after radial strength
of the stent declines and can no longer support the vessel walls.
Additionally, it is expected that the fibrous cap could maintain
the increased radius of curvature after the stent loses mechanical
integrity and after the stent erodes away from the section of the
vessel.
[0139] The stent for treating vulnerable plaque can be
self-expanded or balloon expanded. The stent can include a soft or
elastomeric outer layer on at least the surface that contacts the
plaque region. This outer layer can act as a cushion to prevent
rupture of the vulnerable plaque upon deployment of the stent. In
addition, the deployment pressure of the balloon used to deploy the
stent can be between 30 and 80 psi to prevent rupture of the
vulnerable plaque.
[0140] The fabrication methods of a bioabsorbable stent for use in
the methods of treatment described herein can include the following
steps:
[0141] (1) forming a polymeric tube using extrusion,
[0142] (2) radially deforming the formed tube,
[0143] (3) forming a stent scaffolding from the deformed tube by
laser machining a stent pattern in the deformed tube with an
ultra-short pulse laser,
[0144] (4) forming a therapeutic coating over the scaffolding,
[0145] (5) crimping the stent over a delivery balloon, and
[0146] (6) sterilization with e-beam radiation.
[0147] The stent scaffolding is formed from a semicrystalline
polymer. In particular, a semicrystalline polymer is selected that
has a Tg that is greater than body temperature (about 37.degree.
C.) so that the scaffolding is in a rigid state after implantation
which allows the scaffolding to provide support without excessive
recoil.
[0148] As indicated above, the mechanical properties of a polymer
can be modified by applying stress to a polymer. In particular, the
strength of a polymer can be increased along the direction of the
applied stress. Without being limited by theory, the application of
stress induces molecular orientation along the direction of stress
which increases the strength. Molecular orientation refers to the
relative orientation of polymer chains along a longitudinal or
covalent axis of the polymer chains.
[0149] The fabrication of the polymeric stent includes radially
deforming an extruded polymeric tube about its cylindrical axis.
Radial deformation increases the radial strength of the tubing, and
the subsequently a stent fabricated from the deformed tube. The
increase in strength is due to the induced polymer orientation in
the circumferential direction. It has also been observed that the
deformation increases the fracture toughness of a stent. Both the
increase in radial strength and fracture toughness are important to
the ability of the stent to heal a diseased segment of a blood
vessel.
[0150] Additionally, the stent can have a biaxially oriented
polymer structure. To achieve this, the tube is axially deformed to
provide increased strength in the axial direction, in addition to
being radially deformed. For example, the tube may be axially
deformed by applying a tensile force to the tube along its
cylindrical axis. In some instances, only sufficient tension is
applied to maintain the length of the tube as it is expanded.
[0151] It is generally desirable to deform the tube at a
temperature above the Tg of the polymer. For an exemplary polymer,
PLLA, which has a Tg of about 60.degree. C., the polymer can be
heated to a temperature between 65-120.degree. C. during
deformation. Deforming at such low temperatures favors a high
nucleation density and smaller crystals, which provides high
fracture toughness. The high density of crystallites that are
formed behave a crosslink points that inhibit crack formation and
propagation.
[0152] The polymeric tube is radially deformed using blow molding
through the use of a balloon blower adapted to radially expand a
polymer tube. FIGS. 6A-B illustrate an embodiment of deforming a
polymeric tube. FIG. 6A depicts an axial cross-section of a
polymeric tube 150 with an outside diameter 155 positioned within a
mold 160. Mold 160 limits the radial deformation of polymeric tube
150 to a diameter 165, the inside diameter of mold 160. Polymer
tube 150 may be closed at a distal end 170. Distal end 170 may be
open in subsequent manufacturing steps. A fluid is conveyed, as
indicated by an arrow 175, into an open proximal end 180 of
polymeric tube 150 to increase the pressure inside of the tube. A
tensile force 195 is applied at proximal end 180 and a distal end
170 to axially deform tube 150.
[0153] Polymeric tube 150 is heated by a nozzle directing a heated
gas onto the mold surface. For example, the nozzle can heat the
tube as it translates along the length of the tube. The increase in
pressure inside of polymer tube 150 facilitated by an increase in
temperature of the polymeric tube causes radial deformation of
polymer tube 150 as the nozzle translates, as indicated by an arrow
185. FIG. 6B depicts polymeric tube 150 in a deformed state with an
outside diameter 190 within annular member 160.
[0154] The tube is expanded to a target diameter. The stent pattern
can be cut into the tube with laser machining at the target
diameter. The target diameter can also correspond to the diameter
of a stent prior to crimping.
[0155] The degree of radial deformation may be quantified by
percent radial expansion:
[ Outside Diameter of Deformed Tube Original outside Diameter of
Tube - 1 ] .times. 100 % ##EQU00001##
In some embodiments, percent radial expansion can be 200-500%. In
an exemplary embodiment, the percent radial expansion is about
300%. Similarly, the degree of axial deformation may be quantified
by the percent axial elongation:
[ Length of Deformed Tube Original Length of Tube - 1 ] .times. 100
% ##EQU00002##
The percent axial elongation can be 20-100%.
[0156] Axial polymer orientation is also imparted to a tube during
formation of the tube as the polymer is drawn out of a die during
the extrusion process. The degree of axial orientation of a polymer
provided by the draw down process is related the axial drawn down
ratio:
Inside Diameter of Die Original Inside Diameter of Tube .
##EQU00003##
In an exemplary embodiment the axial drawn down ratio is 2:1 to
7:1.
[0157] The stent pattern is formed in the tube with an
ultrashort-pulse laser. "Ultrashort-pulse lasers" refer to lasers
having pulses with pulse durations shorter than about a picosecond
(=10.sup.-12). Ultrashort-pulse lasers can include both picosecond
and femtosecond (=10.sup.-15) lasers. The stent pattern is formed
with a laser with a pulse width less than 200 fs. In an exemplary
embodiment, the pulse width used is 120 fs. The use of a
femtosecond laser reduces or eliminates damage to polymer material
that is uncut and forms the structure of the stent scaffolding.
[0158] FIG. 7 depicts an exemplary stent pattern 200. Stent pattern
200 can be cut from a polymeric tube using the laser machining
methods described above. Stent pattern 200 is shown in a flattened
condition so that the pattern can be clearly viewed. When the
flattened portion of stent pattern 200 is in a cylindrical form, it
forms a radially expandable stent.
[0159] As depicted in FIG. 7, stent pattern 200 includes a
plurality of cylindrical rings 202 with each ring made up of a
plurality of diamond shaped cells 204. Stent pattern 200 can have
any number of rings 202 depending a desired length of a stent. For
reference, line A-A represents the longitudinal axis of a stent
using the pattern depicted in FIG. 7. Diamond shaped cells 204 are
made up of bar arms 206 and 208 that form a curved element and bar
arms 210 and 212 that form an opposing curved element.
[0160] Pattern 200 further includes linking arms 216 that connect
adjacent cylindrical rings. Linking arms 216 are parallel to line
A-A and connect adjacent rings between intersection 218 of
cylindrically adjacent diamond-shaped elements 204 of one ring and
intersection 218 of cylindrically adjacent diamond shaped elements
204 of an adjacent ring. As shown, linking elements connect every
other intersection along the circumference. Pattern 200 includes
pairs of holes 224 in struts at both ends of the stent to
accommodate radiopaque markers.
[0161] As discussed above, prior to delivery into the body a stent
is compressed or crimped onto a catheter so that it can be inserted
into small vessels. Once the stent is delivered to the treatment
site, it can be expanded or deployed at a treatment site.
Generally, stent crimping is the act of affixing the stent to the
delivery catheter or delivery balloon so that it remains affixed to
the catheter or balloon until the physician desires to deliver the
stent at a treatment site. There are numerous crimpers available
for crimping stents including, but not limited to, the roll
crimper, collet crimper, and wedge crimper.
[0162] The bioabsorable stent is heated and crimped above ambient
temperature. Heating a stent during crimping can reduce or
eliminate radially outward recoiling of a crimped stent which can
result in an unacceptable profile for delivery. In an exemplary
embodiment, a bioabsorbable stent is crimped at a temperature
between 25 and 50.degree. C.
[0163] A crimping device can apply pressure and heat
simultaneously. In these or other embodiments, after crimping, the
crimping device can hold the stent at an elevated temperature,
which may be selected such that it is greater than, equal to, or
less than the selected crimping temperature or may be selected to
specifically exclude temperatures greater than, equal to, or less
than the selected crimping temperature. In some embodiments, the
device crimps the polymeric stent while the stent is heated by
other means.
[0164] The crimped stent is further packaged and sterilized. The
stent is sterilized through exposure to an electron beam (e-beam).
The range of exposure is between 25 and 30 kGy. The radiation
exposure causes degradation in the polymer, particularly the
molecular weight. As discussed above, the radial strength,
mechanical integrity, and erosion profiles are influenced by the
molecular weight. To reduce this degradation, the stent is
sterilized after reducing its temperature below 0.degree. C. by,
for example, placing the stent in a freezer. Additionally, the
initial molecular weight and dose are selected to obtain the
necessary molecular weight for proper functioning of the stent.
[0165] For a stent with an exemplary PLLA scaffolding, the number
average (Mn) and the weight average molecular weight (Mw) of the
scaffolding before and after e-beam sterilization are given in
Table 4.
TABLE-US-00004 TABLE 4 Molecular weight of resin and scaffolding.
Mn (kg/mol) Mw (kg/mol) PLLA Resin 250-300 500-600 PLLA scaffolding
after e-beam 80-100 150-200
[0166] The manufacturing process of a bioabsorbable polymer stent
additionally described in U.S. patent application Ser. No.
11/443,94 which as been published as U.S. Patent Publication No.
20070283552, and is incorporated by reference herein.
Examples
[0167] Some embodiments of the present invention are illustrated by
the following examples and clinical trial information. The examples
and clinical trial information are being given by way of
illustration only and not by way of limitation. The parameters and
data are not be construed to unduly limit the scope of the
embodiments of the invention.
Examples: Embodiments of a Stent can be Fabricated from
Poly(L-Lactide) (PLLA)
Step 1: Tube Manufacturing
[0168] The resin for input into the extruder is granular. An
exemplary PLLA resin has about a 70% crystallinity and Mn=265K,
Mw=520K. Pre-extrusion processing includes baking in a vacuum oven
that removes moisture and residual solvent, both of which can
adversely affect the degradation profiles of the stent. For
instance, moisture can accelerate degradation. The Tm of resin is
about 176.degree. C.
[0169] The resin is extruded in a 1'' single screw extruder used to
form tubing. The parameters are: [0170] Extruded at
420F+/-10.degree. F. (215.degree. C.) [0171] Residence time:
approximately 10 min [0172] Quenched in room temperature water bath
[0173] Die/quench distance is 3/4'' [0174] Pull rate=16 ft/min
[0175] Barrel pressure=2000 psi [0176] Draw down ratio approx 3:1
(ID die to ID of drawn tube) The post-extrusion Mn=180 K, Mw=380K
and crystallinity is 10-15%.
Step 2: Radial Expansion
[0177] The extruded tubing is expanded from 0.018 in inside
diameter (ID)/0.056 in outside diameter (OD) to (0.065 in to 0.080
in ID)/(0.077 in to 0.092 in OD), with 30-80% longitudinal stretch
of the tube. The tubing is expanded by blow molding in a glass
mold. The degree of crystallinity after expansion can be between
30% and 55%. The temperature of the tube during radial expansion
can be between 160.degree. F. and 210.degree. F.
Step 3: Laser Machining and Stent Pattern
[0178] Laser machining is performed with a laser having a 120
femtosecond pulse. The wavelength of the laser is 800 nm.
[0179] Stent struts can have a rectangular or square cross-section.
For example, the struts can measure 0.0065.times.0.0065 in
(150.times.150 micron).
Step 4: Crimping
[0180] The stent is crimped from the cut diameter to a desired
diameter onto a support element, such as a balloon. A sliding wedge
style crimper can be used. The crimp cycle may be between about 30
and 300 seconds. The stent can be heated to a temperature between
28.degree. C. and 48.degree. C. during crimping. The stent can be
crimped from a 0.084 in OD to a 0.053 in OD.
Step 5: Sterilization
[0181] The stent is sterilized by e-beam sterilization with a range
of exposure between 25 and 30 kGy.
Step 5: Deployment
[0182] The crimped stent can be deployed with an outward radial
pressure in the balloon of 7 atm to 0.118 (3.0 mm) ID or a pressure
of up to 16 atm to 0.138 in (3.5 mm) ID.
Clinical Trial Data
[0183] Clinical trials involving implantation of a bioabsorbable
stent in 30 patients were performed. A bioasorbable
everolimus-eluting stent system from BVS of Abbott Vascular, Santa
Clara, Calif. referred to herein as "the BVS stent," was used in
the study. The BVS stent system is made from a bioabsorbable
poly(L-lactide) (PLLA) scaffolding or backbone which is coated with
a more rapidly absorbing poly(D,L-lactide) (PDLLA) layer that
contains and controls the release of the antiproliferative drug,
Everolimus (Novartis, Basel, Switzerland). Clinical trial results
up to 2 years follow-up are reported in, Lancet.com Vol. 373 Mar.
14, 2009, which is incorporated herein by reference.
[0184] The fabrication process of the BVS stent includes the steps
described above, extrusion of a PLLA tube, radial and axial
deformation of the tube, and laser machining a pattern. The PDLLA
coating is applied to the machined backbone prior to crimping on a
delivery balloon.
[0185] FIGS. 8A-B depict images of the BVS stent used in the
studies and is the same as the pattern depicted in FIG. 7. FIG. 8B
depicts a magnified image of the BVS stent. The stent has struts
150 .mu.m thick either directly joined or linked by straight
bridges. Both ends of the stent have two adjacent radiopaque metal
markers. The markers at one end are shown in FIG. 8B.
[0186] The BVS stent shown in FIGS. 8A-B has a backbone of PLLA
coated with PDLLA. As mentioned above, PLLA is a semicrystalline
polymer which is composed of crystalline regions with an amorphous
matrix. The PDLLA in the coating is a random copolymer of D-lactide
and L-lactide. The presence of the D-lactide segments inhibits
crystallization, so the PDLLA is completely amorphous. At room and
physiological temperatures, both PLLA and PDLLA are in the solid
state and below their respective glass transition temperatures. The
PDLLA coating contains and controls the release of the
antiproliferative drug, everolimus. Both PLLA and PDLLA are fully
bioabsorbable. During bioabsorption, the long chains of PLLA and
PDLLA are progressively shortened as ester bonds between repeat
units of lactide are hydrolyzed and small particles less than 2
.mu.m in diameter are phagocytosed by macrophages. Ultimately, PLLA
and PDLLA degrade to lactic acid, which is metabolized via the
Krebs cycle.
[0187] The dose of everolimus on the BVS stent is 98 .mu.g for a 12
mm stent (153 .mu.g for the 18 mm stent). Within 28 days of
implantation, 80% of the drug has eluted from the polymer
coating.
[0188] Some design inputs of the stent are provided in Table 5.
TABLE-US-00005 TABLE 5 Summary of design inputs for BVS stent.
Specification Value Backbone polymer (PLLA) Mw 180,000-200,000 Mn
90,000-100,000 Mass of stent (12 mm length) 5.9 mg Mass/unit length
0.5 mg/mm Crystallinity 45% (as measured by DSC) Strut cross
section 150 micron .times. 150 micron Coating thickness 2 microns
Coating mass 196 .mu.g (1:1 polymer:Everolimus) Coating polymer Mw
66,000 Mn 39,000 Blow molding: Percent radial expansion 300%
Percent axial elongation 50% Laser machining 120 fs laser
[0189] Clinical data up to two years has been obtained. The
techniques and measurements include the following: [0190]
Quantitative coronary angiography (QCA) was used to analyze the
stented segment and the peri-stent segments (defined by a length of
5 mm proximal and distal to the stent edge), as well as their
combination (in-segment analysis); [0191] QCA was used to study
vasomotion at 2 years by measuring mean lumen diameter after
administering either the endothelium independent vasoconstrictor
methylergometrine maleate or the endothelium dependent vasoactive
agent acetylcholine; [0192] Phased array intravascular ultrasound
catheters (IVUS) (EagleEye; Volcano Corporation, Rancho Cordova,
Calif., USA) were used to examine stented vessel segments from
after the procedure and from follow-up; [0193] Optical coherence
tomography (OCT) was used to study strut apposition and changes in
strut appearance at intervention and at 6 months and 2 years after
intervention. [0194] Gray-scale IVUS images were used to assess
appearance of polymer struts; [0195] Gray-scale IVUS was used to
assess echogenecity of polymer struts; [0196] IVUS with virtual
histology (IVUS-VH) (Volcano Corporation, Rancho Cordova, Calif.,
USA) using backscattering of radiofrequency signals was used to
obtain information about tissue composition of the vessel wall;
[0197] Palpography based on IVUS was used to assess deformability
of the vessel wall; [0198] Multislice CT imaging was done 18 months
to determine vessel dimensions;
[0199] The 30 patients had stable, unstable, or silent ischaemia
and a single de-novo lesion that was suitable for treatment with a
single 3.0.times.12 mm or 3.0.times.18 mm stent.
[0200] Some conclusions of analytical techniques are: [0201] IVUS
with echogenicity, IVUS with virtual histology, and OCT indicate
that by 2 years, the BVS stent was incorporated into the vessel
wall and bioabsorbed; [0202] Reduction in molecular weight and mass
had occurred to such an extent that struts were no longer
recognizable by intravascular ultrasound, leaving behind few
visible features, a third of stents were no longer discernible by
OCT; [0203] OCT showed an optically homogeneous vessel wall
structure that, taken together with the documented restoration of
vasomotion, suggests healing of the artery.
[0204] FIG. 9 is a flow chart that summarizes the clinical
population. Four patients were excluded from the per
treatment-evaluable population since they received a non-BVS stent
in addition to the study stent (BVS). The per treatment-evaluable
population was the primary population. Angiographic endpoints,
intravascular ultrasound, and derived morphology parameters were
assessed at 6 months (range 14 days) and at 2 years.
Published information regarding the clinical trials up to 6 months
follow-up can be found in the following: Ormiston et al.,
Lancet.com Vol. 371 Mar. 15, 2008; Tanimoto, S. et al., J. of the
American College of Cardiology, Vol. 52, No. 20, 2008; Ormiston, J.
et al., Catheterization and Cardiovascular Interventions 69:128-131
(2007); Tanimoto, S. et al., Catheterization and Cardiovascular
Interventions 70:515-523 (2007); all of which are incorporated by
reference herein.
[0205] The clinical trials were a single-arm, prospective,
open-label study. Patients were enrolled from four academic
hospitals in Auckland, Rotterdam, Krakow, and Skejby. Patients were
eligible if they were aged 18 years and older and had a diagnosis
of stable, unstable, or silent ischaemia. Additional key
eligibility criteria were the presence of a single, de-novo lesion
in a native coronary artery, which was visually assessed to be less
than 8 mm in length for the 12 mm stent, or less than 14 mm in
length for the 18 mm stent. (18 mm stents were available later
during the enrollment period and were received by only two
patients.) The reference-vessel diameter of the target lesion was
3.0 mm and the stenosis diameter 50% or more and less than 100%,
with a thrombolysis in myocardial infarction (TIMI) flow grade of
more than 1.
Quantitative Coronary Angiography (QCA) Analysis of BVS Stent
[0206] In every patient, the stented segment and peri-stent
segments (defined as 5 mm proximal and distal to the stent edge)
were analyzed by QCA. The following parameters for QCA were
computed: lesion length, minimal luminal diameter (MLD), reference
vessel diameter (RVD), and were obtained by an interpolated method.
Additionally, binary restenosis was computed and is defined in
every segment (stent and peri-stent segment) as diameter stenosis
(DS) of 50% or more at follow-up. Results are presented as paired
matched angiographic views after procedure and at follow-up.
[0207] Table 6 gives baseline characteristics of the per
treatment-evaluable population and intention-to-treat population
including vessel parameters from QCA. Definitions of vessel
characteristics determined from QCA are as follows:
[0208] Post-procedural or post-percutaneous coronary intervention
(PCI) refers to a time point immediately after or almost
immediately after stent deployment.
[0209] "In stent" refers to a stented segment of a vessel.
[0210] "Reference vessel diameter" (RVD) is the diameter of a
vessel in areas adjacent to a diseased section of a vessel that
appear either normal or only minimally diseased.
[0211] "Minimal lumen diameter" (MLD) is the diameter of a diseased
section of a vessel at the site of maximal reduction in the
diameter.
[0212] % "Diameter restenosis" (% DS) is the percent difference
between the reference vessel diameter and the minimal lumen
diameter: (RVD-MLD)/RVD
[0213] "Acute gain" is defined as the difference between pre- and
postprocedural minimal lumen diameter.
[0214] "Late loss" is defined as the difference between minimal
luminal diameter after the procedure or post-percutaneous coronary
intervention (PCI) and minimal luminal diameter at follow-up.
[0215] "Pre-stenting" or "pre-implantation" refers to before
implantation or deployment of the stent at a section of a blood
vessel.
[0216] "Post-stenting" or "post-implantation" refers to a time
shortly after implantation or deployment of the stent at a section
of a blood vessel. Measurements designated post-implantation are
made, for example, immediately after a stent is implanted in a
patient or the same day of implantation.
TABLE-US-00006 TABLE 6 Baseline characteristics of the per
treatment-evaluable population and intention-to-treat population.
Data are mean (SD-standard deviation) or number (%), unless
otherwise indicated. Per treatment-evaluable Intention-to-treat
population population n = 26 n = 30 Age (years) 62 (9) 62 (9) Men
15 (58%) 18 (60%) Current smokers 6 (23%) 6 (20%) Diabetes 1 (4%) 1
(3%) Hypertension needing drugs 16 (62%) 18 (60%) Hyperlipidaemia
needing drugs 16 (62%) 19 (63%) Previous target vessel 2 (8%) 3
(10%) intervention Previous myocardial infarction 1 (4%) 1 (3%)
Stable angina 18 (69%) 21(70%) Unstable angina 7 (27%) 8 (27%)
Silent ischaemia 1 (4%) 1 (3%) Target vessel Left anterior
descending 13 (50%) 15 (47%) Left circumflex 6 (23%) 9 (30%) Right
coronary artery 7 (27%) 7 (23%) AHA/ACC* lesion classification B1
17 (65%) 18 (60%) B2 9 (35%) 12 (40%) Mean diameter of 2 78 (0 47)
2 72 (0 47) reference vessel (mm) Minimum luminal 1.10 (0 26) 1.06
(0 26) diameter (mm) Diameter stenosis (%) 59% (12) 60% (11) Lesion
length (mm) 8 66 (3 97) 9 15 (3 99) *AHA/ACC = American Heart
Association/American College of Cardiology
[0217] Table 7 shows results of the QCA for vessel parameters at
post-PCI, and 6 months and 2 years follow-up (F/U). In-stent
angiographic late loss was 0.48 mm (SD 0.28) at 2 years. The
in-stent late loss is similar to that reported with polymeric
paclitaxel-eluting metallic stents (0.39 mm) [Stone, G. W., et al.,
N Engl J Med 2004; 350: 221-31; Fajadet, J. et al., Circulation
2006; 114: 798-806] more than that with metallic everolimus-eluting
stents (about 0.15 mm) [Grube, E. et al., Circulation 2004; 109:
2168-71; Costa, R A et al.], less than that with a polymeric
zotarolimus-eluting stent (about 0.6 mm) [Meredith I. T., et al.,
Eurolntery 2005; 1: 157-64; Fajadet, J., et al.; Circulation 2006;
114: 798-806] and less than that with bare-metal stents (usually
more than 0.8 mm) [Ormiston et al., Lancet.com Vol. 371 Mar. 15,
2008]. The reference diameter decreased significantly from after
the procedure to 6-month and 2-year follow-up, with an average loss
of about 0.3 mm (Table 7). The late loss in the BVS stent is mostly
due to reduction in stent area, but also induces some intrastent
neointima hyperplasia.
TABLE-US-00007 TABLE 7 Unpaired QCA results for BVS stent.
Difference after Difference Difference after procedure vs 6 months
vs procedure vs p value after p value p value after After 6 2 6
months 2 years 2 years procedure vs 6 months procedure vs QCA
(unpaired) procedure months years (95% CI) (95% CI) (95% CI) 6
months vs 2 years 2 years n 26 26 19 -- -- -- -- -- -- In-stent RVD
(mm) 2.79 2.64 2.43 -0.15 (-0.25 -0.12 (-0.21 -0.29 (-0.40 0.0094
0.0058 <0.0001 (0.41) (0.44) (0.33) to -0.05) to -0.03) to
-0.19) In-stent MLD (mm) 2.32 1.89 1.76 -0.43 (-0.58 -0.08 (-0.19
-0.48 (-0.61 <0.0001 0.23 <0.0001 (0.31) (0.31) (0.35) to
-0.28) to 0.04) to -0.35) In-stent DS (%) 16% 27% 27% 10.51% (5.11
0.58% (-3.57 10.10% (4.58 0.0002 0.81 0.0021 (6) (14) (11) to
15.92) to 4.72) to 15.62) In-stent late -- 0.43 0.48 -- 0.08 (-0.04
-- -- 0.233 -- loss (mm) (0.37) (0.28) to 0.19) Proximal late --
0.23 0.34 -- 0.11 (-0.01 -- -- 0.0553 -- loss (mm) (0.31) (0.33) to
0.23) Distal late -- 0.23 0.36 -- 0.16 (0.04 -- -- 0.0091 -- loss
(mm) (0.27) (0.37) to 0.29) In-stent absolute 4.81 3.22 2.68 -1.59
(-2.67 -0.02 (-0.45 -1.89 (-2.65 0.0002 0.93 <0.0001 minimal
luminal (1.75) (1.93) (1.21) to -0.52) to 0.41) to -1.12) area ED
(mm2) In-stent minimal 5.53 3.86 3.18 -1.67 (-2.92 -0.15 (-0.79
-2.04 (-2.96 <0.0001 0.73 <0.0001 luminal cross (2.11) (2.26)
(1.55) to -0.42) to 0.49) to -1.11) sectional area VD (mm2)
In-segment -- 0.35 0.37 -- 0.07 (-0.06 -- -- 0.42 -- late loss (mm)
(0.32) (0.27) to 0.20) In-stent binary -- 7.7% 0 -- -5.3% (-25.2 --
-- 1 -- restenosis (%) (2/26) (0/19) to 13.8) In-segment binary --
7.7% 0 -- -5.3% (-25.2 -- -- 1 -- restenosis (%) (2/26) (0/19) to
13.8)
[0218] As shown in Table 7, between 6 months and 2 years, there was
no significant differences in in-stent minimal lumen diameter,
percentage of diameter stenosis, and in-stent late loss. Therefore,
the significant decrease in minimal lumen diameter, reference
vessel diameter, and luminal area already recorded at 6 months,
remained significant at 2 years. Table 7 shows distal segment late
loss increased significantly at 6 months and 2 years.
[0219] Acute stent recoil measured by QCA immediately after stent
deployment was slightly higher than that of a matched population
with lesions, who were receiving an everolimus-eluting metallic
stent of 3 mm in diameter. Tanimoto, S., et al., Catheter
Cardiovasc Intery 2007; 70: 515-23. The acute recoil is on average
6.9% for the BVS stent and 4.3% for the everolimus-eluting metallic
stent. The acute recoil of BVS stent in vessels less than 3 mm is
8.4% and 11.8% in a calcified lesion.
Explain statement from Lancet paper--Data that shows this: Patency
and absence of binary restenosis was established non-invasively,
and subsequently confirmed by conventional angiography.
Quantitative IVUS Measurements of Vessel
[0220] Stented vessel segments after the procedure and at follow-up
were examined with phased array intravascular ultrasound (IVUS)
catheters (EagleEye, Volcano Corporation, Rancho Cordova, Calif.,
USA) with automated pullback at 0.5 mm per second. The region
beginning 5 mm distal to and extending 5 mm proximal to the stented
segment was examined. The vessel area and mean lumen area were
measured with a computer-based contour detection program (Curad,
version 3.1).
[0221] As shown in Table 8 and FIGS. 10 and 11 (described in more
detail below), gray-scale IVUS showed significant increase in
minimal luminal area and average luminal area and volume, together
with a significant decrease in plaque area and volume between 6
months and 2 years. With the exception of the minimal luminal area,
findings for vessel area, average luminal area, plaque area, and
lumen area stenosis at 2 years did not differ significantly from
the measurement taken immediately after the procedure (table 3).
The vessel area and volume remained constant between the
follow-ups, showing the absence of significant remodeling.
TABLE-US-00008 TABLE 8 Clinical trial IVUS analysis. Difference
after Difference Difference after procedure vs 6 months vs
procedure vs p value after p value p value after After 6 2 6 months
2 years 2 years procedure vs 6 months procedure vs procedure months
years (95% CI) (95% CI) (95% CI) 6 months vs 2 years 2 years n 25
25 19 -- -- -- -- -- -- Vessel (EEM) 13.9 13.79 12.75 -0.06 (-0.49
-0.19 (-0.98 -0.21 (-1.21 0.98 0.24 0.68 area (mm2) (3.74).dagger.
(3.84) (3.43) to 0.37) to 0.59) to 0.78) Vessel volume 173.17
187.65 178.21 12.29 (-4.11 1.01 (-27.27 4.36 (-22.03 0.21 0.86 0.71
(mm3) (52.04).dagger. (72.75) (64.63) .dagger-dbl. to 28.69) to
29.30) to 30.76) Average lumen 6.04 5.19 5.47 -1.01 (-1.30 0.68
(0.04 -0.40 (-1.18 <0.0001 0.0174 0.12 area (mm2) (1.12) (1.33)
(2.11) to -0.71) to 1.32) to 0.38) Lumen volume 78.23 70.66 77.60
-9.20 (-16.84 12.42 (-1.19 -1.33 (-16.29 0.0032 0.0443 0.97 (mm3)
(22.98) (26.88) (35.98) .dagger-dbl. to -1.56) to 26.03) to 13.62)
Plaque area 7.44 8.60 7.10 0.93 (0.45 -1.06 (-1.48 0.01 (-0.71
<0.0001 0.0001 0.80 (mm2) (2.83).dagger. (2.85) (2.02) to 1.40)
to -0.64) to 0.72) Plaque volume 94.56 116.99 98.75 21.11 (9.51
-13.38 (-30.22 4.09 (-11.81 <0.0001 0.0063 0.71 (mm3)
(35.43).dagger. (48.96) (36.47) .dagger-dbl. to 32.72) to 3.47) to
20.00) Minimal lumen 5.09 3.92 4.34 -1.26 (-1.55 0.76 (0.22 -0.59
(-1.26 <0.0001 0.0026 0.0323 area (mm2) (1.02) (0.98) (1.74) to
-0.96) to 1.31) to 0.08) Lumen area 15.83% 23.62% 20.38% 7.28%
(3.54 -4.12% (-8.30 4.07% (-1.30 0.0009 0.0569 0.0799 stenosis (%)
(7.64) (10.25) (6.92) to 11.02) to 0.07) to 9.44) Projected MLD
2.28 2.04 2.17 -0.26 (-0.35 0.19 (0.06 -0.07 (-0.26 <0.0001
0.0052 0.23 (mm) (0.26) (0.26) (0.43) to -0.18) to 0.33) to 0.11)
p-values per Wilcoxon's signed rank test
Serial Assessment of IVUS-VH Results
[0222] FIG. 10 and Table 9 provide serial assessments of IVUS-VH.
FIG. 10 depicts gray-scale IVUS-VH images (top) and corresponding
radiofrequency processed images (bottom) of a vessel of a patient
before stenting, post-stenting, 6 months after stenting, and 2
years after stenting.
TABLE-US-00009 TABLE 9 Quantitative IVUS-VH results. After
Difference after Difference Difference after procedure procedure vs
6 months vs procedure vs p value after p value p value after IVUS
virtual 6 2 6 months 2 years 2 years procedure vs 6 months
procedure vs histology (unpaired).sctn. months years (95% CI) (95%
CI) (95% CI) 6 months vs 2 years 2 years n 25 25 18 -- -- -- -- --
-- Dense calcium (%) 20.65% 26.42% -8.93% (-13.64 2.81% (-4.08
-5.87% (-13.84 0.0003 0.64 0.21 29.82% (15.57) (11.50) (15.76) to
-4.22) to 9.70) to 2.11) Dense calcium 0.94 0.81 -0.11 (-0.36 -0.11
(-0.40 -0.16 (-0.45 0.5046 0.31 0.21 area (mm2) (0.64) (0.67) to
0.15) to 0.17) to 0.13) 1.02 (0.58) Fibro-fatty 7.19% 5.47% 2.94%
(0.40 -0.41% (-3.50 1.72% (-0.09 0.0142 0.85 0.21 tissue (%) (6.17)
(5.22) to 5.48) to 2.67) to 3.53) 4.31% (3.35) Fibro-fatty 0.40
0.19 0.19 (0.02 -0.12 (-0.27 0.01 (-0.05 0.0096 0.0267 0.80 area
(mm2) (0.43) (0.24) to 0.35) to 0.04) to 0.07) 0.21 (0.22) Fibrous
(%) 50.54% 43.66% 11.79% (6.84 -3.38% (-10.59 6.75% (-0.68
<0.0001 0.35 0.20 38.83% (13.41) (12.69) (14.69) to 16.74) to
3.82) to 14.17) Fibrous area (mm2) 2.62 1.35 0.80 (0.51 -0.92
(-1.25 -0.25 (-0.61 <0.0001 <0.0001 0.25 1.72 (1.22) (1.44)
(0.92) to 1.10) to -0.58) to 0.10) Necrotic core (%) 21.62% 24.45%
-5.79% (-9.45 0.99% (-2.48 -2.60% (-5.83 0.0028 0.64 0.30 27.04%
(7.00) (8.70) (6.84) to -2.13) to 4.46) to 0.63) Necrotic core 1.13
0.79 -0.13 (-0.38 -0.26 (-0.53 -0.28 (-0.50 0.342 0.1089 0.0268
area (mm2) (0.87) (0.51) to 0.11) to -0.00) to -0.05) 1.17
(0.82)
[0223] Backscattering of radiofrequency signals provides
information about tissue composition of the vessel wall (with use
of IVUS-VH). In the radiofrequency processed images, typically the
necrotic core is represented as red areas on ultrasound
cross-sections, dense calcium as white, fibro-fatty tissue as
yellow-green, and fibrous tissue as green, and expressed as
percentages (per cross-section, with the total equaling 100%). On
every cross-section, polymeric stent struts were detected as areas
of apparent dense calcium due to the strong backscattering
properties of the polymer.
[0224] The change in quantitative analysis of these areas between
implantation and follow-up is used as a surrogate assessment of the
polymer bioabsorption process. Therefore, a substantial portion of
the 29.82% at post-stenting measured as dense calcium is polymeric
stent struts. The white areas in the post-stenting radiofrequency
images of FIG. 10 represent the stent struts. At 6 month follow-up,
the dense calcium measurement has decreased to 20.65%, indicating
that the struts are partially absorbed at this time point. The
lumen area is shown in each radiofrequency image.
[0225] As shown, the lumen area increases from 3.9 mm.sup.2 at
pre-stenting to 7.1 mm.sup.2 at post-stenting, illustrating the
enlargement of the lumen by the stent deployment. The increase in
lumen diameter is apparent. Although there is a slight decrease in
lumen area to 6.9 mm.sup.2 at 6 months follow-up, the lumen size is
maintained with no apparent restenosis. At 2 years follow up, the
lumen areas has increased to 10.1 mm.sup.2 and the lumen size is
maintained.
[0226] From the OCT measurements discussed below, the stent is
completely or almost completely absorbed at 2 year follow-up. Thus,
the slight increase in the dense calcium measurement of 26.42% at 2
year follow-up in Table 9 is not a measurement of polymer strut,
rather it is likely a measurement of calcification of the void left
by the absorbed stent strut.
[0227] The IVUS-VH assessments showed that the percentage of each
plaque component did not differ significantly between 6 months and
2 years (Table 9, FIG. 10). The absolute fibro-fatty area and
fibrous-plaque area decreased significantly between 6 months and 2
years (Table 9). When compared with measurements taken immediately
after the procedure, none of the 2-year parameters differed
significantly, apart from necrotic core area (Table 9).
[0228] A shown in the lower three panels of FIG. 10, at
pre-stenting there is a large plaque with a necrotic core. The
arrow in the pre-stenting frame indicates the necrotic core is in
contact with the lumen. After stenting, the lumen has enlarged to
7.1 mm.sup.2 and two arrows point at two struts that are depicted
as (pseudo) dense calcium. At 6 months, the two struts remained
visible as indicated by the arrows with partial reduction of the
luminal area. At 2 years, the endoluminal pseudo-dense calcium
depicting struts has disappeared, which is consistent with strut
absorption. There is now a fibrous cap covering the necrotic core,
as indicated by the arrows, and late enlargement of the lumen (10.1
mm.sup.2) and shrinkage of the plaque.
IVUS: Tissue Echogenicity and Palpography
[0229] FIG. 11 depicts IVUS images with tissue echogenicity of
two-dimensional radial slices of an implant site of a single
patient at post-PCI and at 6 months follow-up. Table 10 and FIG. 11
show that a significant reduction in percentage of hyperechogenic
tissue between after the procedure and at 6 months, and between 6
months and 2 years, in the intention-to-treat population was
detected. The residual level of hyperechogenicity at 2 years was
similar to the natural hyperechogenicity of plaques (7.7% [SD 6.5]
vs 5.6% [4.8], n=12) that was measured in one investigating center
(Thorax Center, Rotterdam Netherlands). Importantly, persistent and
late acquired incomplete apposition detected at 6 months and
previously reported, 6 were longer detectable at 2-year
follow-up.
TABLE-US-00010 TABLE 10 Tissue echogenicity and palpography
results. Difference after Difference Difference after procedure vs
6 months vs procedure vs p value after p value p value after After
6 2 6 months 2 years 2 years procedure vs 6 months procedure vs
procedure months years (95% CI) (95% CI) (95% CI) 6 months vs 2
years 2 years Palpography (unpaired) n 24 23 17 -- -- -- -- -- --
Strain values 0.16 0.28 0.31 0.12 (0.06 0.02 (-0.05 0.13 (0.06
0.0002 0.81 0.0052 (0.10) (0.12) (0.17) to 0.17) to 0.08) to 0.21)
Echogenicity (ITT) n 27 26 21 -- -- -- -- -- -- Hyperechogenicity
(%) 18.5% 10.3% 7.7% -8.15% (-11.00 -3.75% (-6.20 -12.81% (-16.19
<0.0001 0.001 <0.0001 (9.1) (7.6) (6.5) to -5.31) to -1.29)
to -9.44)
[0230] In FIG. 11, the two top images are gray-scale IVUS images at
post-PCI and 6 months follow-up, respectively. The bottom images
are color-coded echogenicity images at post-PCI and 6 months
follow-up, respectively. The green or lighter-colored portions are
hyperechogenic material that correspond to stent struts. The red or
darker areas are hypoechogenic tissue components that do not
include stent struts. At 6 months follow-up, the images show that
the stent struts are less pronounced, indicating absorption of
stent struts at this time point.
[0231] FIG. 12 depicts IVUS three-dimensional images with tissue
echogenicity containing information along the entire axis of the
stented site at post-PCI (left) and at 6 months follow-up (right).
These images further illustrate the extent of absorption of the
struts at 6 months follow-up.
[0232] In the palpography assessments, the underlying principle is
that softer tissue is more readily deformed than is harder or
scaffolded tissue when force (e.g., pulsatile arterial pressure) is
applied. The rationale of this analysis for the study was to detect
some subtle changes in strain resulting from scaffolding and late
bioabsorption of the stent. The deformability of vessel wall was
quantified with the analysis of back-scattering radiofrequency
signals at different diastolic pressure levels.
[0233] Table 10 shows that the cumulative strain value increased
significantly from after the procedure to 6-month follow-up, with
no subsequent changes between 6 months and 2 years. The vessel wall
deformability reappears to some extent in the initial 6 months and
remained stable.
Quantitative OCT Measurements of Vessel
[0234] A commercially available OCT system was used in a subgroup
of patients. This technique, with use of an infrared light source,
has a resolution of 15 .mu.m which is about ten times higher than
that of intravascular ultrasound and therefore allows visualization
of intracoronary structures in great detail. The light source is a
1310 nm broadband super luminescent diode with an imaging depth of
about 1.5 mm, an axial resolution of 15 and a lateral resolution of
25 The imaging probe (ImageWire LightLab Imaging Inc, Westford,
Mass., USA) has a maximum outer diameter of 0.4826 mm
(0.cndot.019'') and contains a 0.1524 mm (0.cndot.006'')
fiber-optic imaging core and a distal radiopaque spring tip, which
is similar to conventional guide wires. An OCT catheter (Helios
proximal occlusion catheter) is initially advanced distal to the
area of interest over a conventional coronary guide wire, which is
then replaced with the OCT imaging wire (ImageWire).
[0235] OCT data for the stent length, lumen area, minimal lumen
area (MLA), minimal lumen diameter (MLD), and lumen volume are
shown in Table 11 for PCI, 6 months follow-up, and 2 years
follow-up. OCT data show a decrease in lumen area mean, MLA, and
MLD at 6 months follow-up. Between 6 months and 2 years, there is
an increase in mean and minimal lumen area and luminal volume.
TABLE-US-00011 TABLE 11 Clinical trial OCT analysis. Difference
after Difference Difference after procedure vs 6 months vs
procedure vs p value after p value p value after After procedure 6
2 6 months 2 years 2 years procedure vs 6 months procedure vs
Optical CT (serial) months years (95% I) (95% CI) (95% CI) 6 months
vs 2 years 2 years n 7 7 7 -- -- -- -- -- -- Discernible struts 403
368 264 -- -- -- -- -- -- Mean lumen area 6.53 4.72 5.80 -1.81
(-3.39 1.08 (-0.93 -0.74 (-3.42 0.0313 0.22 0.38 (mm2) (0.91)
(1.13) (2.93) to -0.23) to 3.08) to 1.94) Minimal lumen 4.50 2.65
3.80 -1.84 (-3.48 1.15 (-0.14 -0.7 (-2.62 0.0156 0.0781 0.47 area
(mm2) (1.03) (1.49) (2.42) to -0.21) to 2.43) to 1.22) Mean lumen
2.87 2.41 2.63 -0.46 (-0.87 0.23 (-0.17 -0.23 (-0.79 0.0313 0.22
0.30 diameter (mm) (0.21) (0.31) (0.61) to -0.05) to 0.62) to 0.33)
Lumen volume 84.1 58.0 74.1 (mm3) Stent length 12.7 12.5 12.7
(mm)
[0236] OCT imaging data (discussed below) show absorption of the
stent into artery walls and that the blood vessel lining of
arteries treated with the stent looks more uniform after two years
than it did immediately post-treatment.
[0237] The significant increase in the average luminal area and
minimal luminal area measured by IVUS and OCT between 6 months and
2. years contrasts with the nonsignificant decrease in angiographic
luminal dimensions during that period.
Several explanations for the discordant late luminal changes
between angiography and intracoronary imaging have been considered.
Lancet.com Vol. 373 Mar. 14, 2009, p. 907. Stent Strut Apposition
and Appearance from OCT
[0238] The appearance of struts from OCT images changes at
follow-up. The strut appearance can be characterized into four
groups: [0239] A "preserved box" has sharp defined, bright
reflection borders with preserved box shaped appearance. The strut
body shows low reflection. A preserved box is an image of a strut
which has undergone little or no change. [0240] An "open box" has
luminal and abluminal "long-axis" borders thickened and bright
reflection. "Short axis" borders are not visible. An open box is an
image of a strut which has partially started to be dissolved.
[0241] A "dissolved bright box" is a partially visible bright spot
with contours poorly defined and no box shaped appearance. A
dissolved bright box is an image of a strut that has mostly
dissolved. [0242] A "dissolved black box" is a black spot with
contours poorly defined and often confluent and no box shaped
appearance. A dissolved black box is an image of a strut that has
mostly dissolved and all or part of the void replaced by inorganic
material.
[0243] Seven patients had serial data for OCT immediately after the
procedure, at 6 months, and at 2 years (intention-to-treat
population). The number of apparent struts decreased from 403 at
baseline to 368 at 6-month follow-up and to 264 at 2 years (34.5%
reduction over 2 years. For preserved box, appearance of stent
strut changed from n=0 at 6 months to n=9 at 2 years, for open box
from n=143 to n=68, for dissolved bright box from n=225 to n=185,
and for dissolved black box from n=56 to n=25.
[0244] FIGS. 13A-B depicts serial assessment of stent struts by
OCT. FIG. 13A shows, after stenting, incomplete apposition of
struts (preserved box) in front of a side-branch ostium, At 6
months, persistent incomplete stent apposition (arrow) and resolved
incomplete stent apposition (arrowhead), with open box appearance.
At 2 years, there is now smooth appearance of the endoluminal
lining without strut malapposition since struts have been absorbed.
There is guidewire shadowing (at the top of the image), and a strut
is still just discernible as a bright spot (arrow).
[0245] In FIG. 13B, complete apposition of strut (box appearance)
after the procedure is shown. At 6 months, there is late acquired
incomplete stent apposition of the struts (preserved box
appearance) with tissue bridging connecting the struts (arrow). The
endoluminal lining is corrugated. At 2 years, the smooth
endothelial lining with almost circular cross section. Generally,
the struts are no longer discernible, although there is a bright
reflection that could indicate a strut (arrow). Asterisk indicates
a side branch.
[0246] All apparent struts were well covered and apposed to the
vessel wall. All incomplete appositions (incomplete, persistent,
and late acquired incomplete stent apposition) were resolved. The
lumen shape was regular with smooth, well delineated borders in all
cases, and we recorded no intraluminal tissue. The coronary vessel
wall showed a homogenous, bright backscattering appearance, with no
signs of tissue optical heterogeneity.
[0247] FIGS. 14A-D depict exemplary OCT images at 6 months
follow-up of a preserved box, open box, bright dissolved box, and
black dissolved box, respectively. At 6 months follow-up, 18 (3%)
of 671 struts had a preserved box, 203 (30%) had an open box, 332
(50%) had a dissolved bright box, and 118 (18%) had a dissolved
black-box appearance.
[0248] FIGS. 14C-D indicate that the stent has started to erode at
6 months, which means that the stent has started to lose mechanical
integrity. The erosion and loss of mechanical integrity suggests
that the stent has lost radial strength well before 6 months. If
the stent had maintained mechanical integrity, all regions would
look like FIG. 14A.
[0249] FIGS. 15A-C depict complete OCT images post-PCI, at 6 months
follow-up, and 2 years follow-up, respectively. The dark
rectangular areas indicated in FIG. 15A correspond to the struts of
the stent. In FIG. 15B, the dark areas are replaced by bright areas
which indicate partially dissolved struts. No struts are apparent
in FIG. 15C which suggest that the struts are almost or completely
dissolved.
Bioabsorption as Shown by IVUS, IVUS-VH, and OCT for 2 Years
Follow-Up
[0250] FIGS. 16A-B, 17A-B, and 18A-B are IVUS IVUS-VH, and OCT
images, respectively, for one patient post-PCI and at 2 years
follow-up. The top three frames from left to right, 15A, 16A, and
17A, are IVUS, VH, and OCT images post-PCI. The bottom three frames
from left to right, 16B, 17B, and 18B, are IVUS, VH, and OCT images
at 2 years.
[0251] In FIG. 16A, the white spots on the inner side of the vessel
that are indicated correspond to stent struts. As shown in FIG.
16B, the whites spots are not present at 2 years follow-up.
[0252] In FIG. 17A-B, the white spots correspond to high density
material and the stent struts are apparent around the vessel as
indicated. As shown in FIG. 17B, after 2 years most of the stent
struts have disappeared. The two white spots indicated may not be
stent struts, but may be mineralization.
[0253] In FIG. 18A, black spots as indicated are stent struts that
are in the inner side of the vessel. As shown in FIG. 18B, the
black spots are replaced by bright yellow regions, as indicated,
which indicates that stent struts are gone.
Endothelialization
[0254] Tissue coverage was present in 664 (99%) of the struts seen
with an OCT image of a patient at 6 months (only seven had no
tissue coverage as detected by OCT). FIG. 18A depicts an OCT image
of a section in which arrows indicate complete tissue coverage of a
strut. FIG. 18B depicts an OCT section in which arrows indicate
incomplete tissue coverage of a strut.
Restoration of Vasomotion
[0255] Previously stented portion of arteries demonstrated the
ability to expand and contract in a manner similar to a vessel that
has never been stented. To study vasomotion at 2 years, either the
endothelium independent vasoconstrictor methylergometrine maleate
(methergin, Novartis, Basel, Switzerland), or the endothelium
dependent vasoactive agent acetylcholine (Ovisot, Daiichi-Sankyo,
Tokyo, Japan) was given. Potential restoration of unstented artery
movement to coronary blood vessel after the bioabsorbable stent was
absorbed was revealed at two years with the drugs acetylcholine and
nitroglycerin used in nine patients and methergine in seven
patients. Acetylcholine and nitroglycerin tend to induce
vasodilation and methergine tends to induce vasoconstriction in
blood vessels.
[0256] A method of subsegmental analysis was used to calculate the
mean lumen diameter for the stented segment and its adjacent
segments 5 mm proximally and distally. Each segment was divided
into several subsegments and a mean lumen diameter of each segment
was calculated from the subsegments. Measurements of the luminal
diameter are shown for the stented segment, proximal segment, and
the distal segment.
[0257] FIG. 20A shows angiography measurements for patients treated
with methergine and nitroglycerine. For each patient, the luminal
diameter was measured pre-methergine treatment, with methergine
treatment, and with nitroglycerine treatment. There was significant
vasoconstriction in proximal and stented segments. After
nitroglycerin, the three segments (proximal, stented, and distal)
dilated significantly with their diameters returning to their
baseline values.
[0258] FIG. 20B shows angiography measurements for patients treated
with acetylcholine and nitroglycerin. Five patients had
vasodilation of at least 3% in mean luminal diameter. Nitrates
induced a significant vasodilatation in the stented and distal
segments. These results show the restoration of vasomotor function
in the stented segment.
[0259] Table 12 provides vessel measurements for one patient at 2
years that demonstrate vasoconstriction and vasocompression. The
reappearance of vasomotion of the stented and persistent segments
in response to methergine or acetylcholine suggests that vessel
vasoreactivity has been restored and that a physiological response
to vasoactive stimulus might occur anew. Unlike previous studies
reporting endothelial dysfunction in the distal segment, this study
shows vasomotor tone in the stented segment, once the scaffolding
properties of the stent had disappeared as a result of its
bioabsorption. Five of the nine patients tested with acetylcholine
showed vasodilatation (at least 3% of the mean diameter) during the
highest dose infused. This suggests direct vasodilator or
flow-mediated response to acetylcholine and thus the presence of
functionally active endothelium at the site of the stent
implantation.
TABLE-US-00012 TABLE 12 Vessel diameters pre-and post-methergine
and post nitroglycerine treatment at 2 years follow-up for one
patient. Pre- Post Post Mether- Methergine Change Nitro- Change
gine (5 min) N (%) glycerine N (%) In-Stent 2.62 mm 1.98 mm -0.64
mm 2.58 mm +0.60 mm Mean (-24.4%) (+30.3%) Diameter Mean 2.70 mm
2.08 mm -0.62 mm 2.62 mm +0.54 mm Diameter 2-17 (-23.0%) (+26.0%)
mm Distal to Stent
Multi-Slice CT Results
[0260] Multi-slice CT imaging was done 18 months after the index
procedure. Single or dual-source, 64-slice spiral CT with
intravenous contrast enhancement and electrocardiograph-gated image
reconstruction was done (Siemens Definition, Forchheim, Germany
[n=18]; GE Lightspeed, Milwaukee, USA [n=5]; Philips Brilliance,
Best, Netherlands [n=2]). The amount of in-stent stenosis was
measured using a semi-automated software program for vessel
segmentation and lumen area quantification (Circulation, Siemens,
Forchheim, Germany).
[0261] The BVS stent is undetected by multi-slice CT, apart from
the platinum markers at each end of the stent. Along the
automatically constructed center-lumen line, the cross-sectional
lumen area was measured at 0.3 mm longitudinal intervals within the
stented segment. The lumen diameter was calculated from the
measured area, assuming a circular shape of the lumen area.
Severity of in-stent diameter stenosis or area stenosis was
calculated as a ratio of the smallest in-stent lumen diameter or
area and the reference vessel diameter or area, which was
calculated by interpolation of the proximal and distal lumen
reference. Additionally, the length of vessel was measured between
the platinum stent markers. The quantitative results in Table 13
show that all stents were qualitatively patent.
TABLE-US-00013 TABLE 13 Quantitative multi-slice CT results. 2
years N 24 Mean luminal area (mm.sup.2) 5 2 (1.3) Minimal luminal
area mm.sup.2) 3 6 (0.9) Reference area (mm.sup.2) 5 5 (1.0) Mean
area stenosis (%) 34% (15) Minimal diameter (mm) 2 12 (0.26) Mean
diameter stenosis (%) 19% (9)
Summary of Cardiac Events for Intent to Treat Population
[0262] Table 14 summarizes the cardiac events of the intent to
treat clinical population. Patients were followed out to two years.
No stent thrombosis and no major adverse cardiac events were
observed. No new MACE events between 6 months and 2 years. No stent
thrombosis was observed up to 2 years.
TABLE-US-00014 TABLE 14 Cardiac events of intent to treat
population. 6 Months 12 Months 18 Months 2 Years 30 29 29 28
Hierarchical Patients Patients** Patients** Patients** Ischemia
Driven 3.3% (1)* 3.4% (1)* 3.4% (1)* 3.6% (1)* MACE (%) Cardiac
Death (%) 0.0% (0) 0.0% (0) 0.0% (0) 0.0% (0) MI (%) 3.3% (1)* 3.4%
(1)* 3.4% (1)* 3.6% (1)* Q-Wave MI 0.0% (0) 0.0% (0) 0.0% (0) 0.0%
(0) Non Q-Wave MI 3.3% (1)* 3.4% (1)* 3.4% (1)* 3.6% (1)* Ischemia
Driven 0.0% (0) 0.0% (0) 0.0% (0) 0.0% (0) TLR (%) by PCI 0.0% (0)
0.0% (0) 0.0% (0) 0.0% (0) by CABG 0.0% (0) 0.0% (0) 0.0% (0) 0.0%
(0) *Same patient--this patient also underwent a TLR, not qualified
as ID-TLR (DS = 42%) **One patient missed the 9, 12, 18 month and 2
year visits. One patient died from a non-cardiac cause 706 days
post-procedure MACE = major adverse cardiac events MI =
post-myocardial infarction (MI) TLR = target lesion
revascularization ID-TLR = ischemia driven target lesion
revascularization PCI = post-percutaneous coronary intervention
CABG = Coronary Artery Bypass Graft
In Vitro Testing of Radial Strength
[0263] The radial strength of BVS stent was tested in vitro at four
time points. The tests show the effect of degradation on radial
strength. The types of stents used in the in vitro tests are the
same as those used in human clinical trials. The results of these
trials are discussed below. The in vitro tests were performed by
immersing stents in a phosphate buffered saline solution simulating
a vascular environment.
[0264] The radial strength of stents in the four arms was tested.
Each arm initially contained 3 stents. The radial strength of the 3
stents was tested at time zero or no exposure to the solution, two
weeks, 28 days, and three months. The radial strength was measured
by a flat plate compression test using a machine obtained from
Instron in Canton, Mass.
[0265] The representative results of the in vitro tests from arm B
stents are shown in FIG. 20. The results indicate a small change in
radial strength between time zero and 28 days. The radial strength
decreased at 28 days and was undetectable at three months.
Therefore, the stent lost radial strength between 28 days and three
months.
[0266] "Radial strength" of a stent is defined as the pressure at
which a stent experiences irrecoverable deformation.
[0267] "Stress" refers to force per unit area, as in the force
acting through a small area within a plane. Stress can be divided
into components, normal and parallel to the plane, called normal
stress and shear stress, respectively. Tensile stress, for example,
is a normal component of stress applied that leads to expansion
(increase in length). In addition, compressive stress is a normal
component of stress applied to materials resulting in their
compaction (decrease in length). Stress may result in deformation
of a material, which refers to a change in length. "Expansion" or
"compression" may be defined as the increase or decrease in length
of a sample of material when the sample is subjected to stress.
[0268] "Strain" refers to the amount of expansion or compression
that occurs in a material at a given stress or load. Strain may be
expressed as a fraction or percentage of the original length, i.e.,
the change in length divided by the original length. Strain,
therefore, is positive for expansion and negative for
compression.
[0269] "Strength" refers to the maximum stress along an axis which
a material will withstand prior to fracture. The ultimate strength
is calculated from the maximum load applied during the test divided
by the original cross-sectional area.
[0270] "Modulus" may be defined as the ratio of a component of
stress or force per unit area applied to a material divided by the
strain along an axis of applied force that result from the applied
force. For example, a material has both a tensile and a compressive
modulus.
[0271] The tensile stress on a material may be increased until it
reaches a "tensile strength" which refers to the maximum tensile
stress which a material will withstand prior to fracture. The
ultimate tensile strength is calculated from the maximum load
applied during a test divided by the original cross-sectional area.
Similarly, "compressive strength" is the capacity of a material to
withstand axially directed pushing forces. When the limit of
compressive strength is reached, a material is crushed.
[0272] The underlying structure or substrate of an implantable
medical device, such as a stent can be completely or at least in
part made from a biodegradable polymer or combination of
biodegradable polymers, a biostable polymer or combination of
biostable polymers, or a combination of biodegradable and biostable
polymers. Additionally, a polymer-based coating for a surface of a
device can be a biodegradable polymer or combination of
biodegradable polymers, a biostable polymer or combination of
biostable polymers, or a combination of biodegradable and biostable
polymers.
[0273] It is understood that after the process of degradation,
erosion, absorption, and/or resorption has been completed, no part
of the stent will remain or in the case of coating applications on
a biostable scaffolding, no polymer will remain on the device. In
some embodiments, very negligible traces or residue may be left
behind. For stents made from a biodegradable polymer, the stent is
intended to remain in the body for a duration of time until its
intended function of, for example, maintaining vascular patency
and/or drug delivery is accomplished.
[0274] While particular embodiments of the present invention have
been shown and described, it will be obvious to those skilled in
the art that changes and modifications can be made without
departing from this invention in its broader aspects. Therefore,
the appended claims are to encompass within their scope all such
changes and modifications as fall within the true spirit and scope
of this invention.
* * * * *