U.S. patent application number 10/165504 was filed with the patent office on 2005-06-02 for cardiac support device.
Invention is credited to Girard, Michael J., Palme, Donald F. II, Shapland, J. Edward.
Application Number | 20050119519 10/165504 |
Document ID | / |
Family ID | 29710450 |
Filed Date | 2005-06-02 |
United States Patent
Application |
20050119519 |
Kind Code |
A9 |
Girard, Michael J. ; et
al. |
June 2, 2005 |
CARDIAC SUPPORT DEVICE
Abstract
A highly compliant and elastic cardiac support device is
provided. The device is constructed from a biocompatible material
is applied to an external surface of a heart. The device can be
used to resist dilatation of the heart, to provide acute wall
support, or to enhance reduction in the size of the heart using
stored potential energy, without interfering with systolic
contraction.
Inventors: |
Girard, Michael J.; (Lino
Lakes, MN) ; Shapland, J. Edward; (Vadnais Heights,
MN) ; Palme, Donald F. II; (Princeton, MN) |
Correspondence
Address: |
ACORN CARDIOVASCULAR, INC.
P.O. BOX 2903
MINNEAPOLIS
MN
55402
US
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Prior
Publication: |
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Document Identifier |
Publication Date |
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US 0229265 A1 |
December 11, 2003 |
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Family ID: |
29710450 |
Appl. No.: |
10/165504 |
Filed: |
June 7, 2002 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10165504 |
Jun 7, 2002 |
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09593251 |
Jun 13, 2000 |
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6482146 |
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Current U.S.
Class: |
600/37 |
Current CPC
Class: |
A61F 2/2481
20130101 |
Class at
Publication: |
600/037 |
International
Class: |
A61F 002/00 |
Claims
What is claimed is:
1. A device for treating diseases of a heart, said device
comprising: a biocompatible material sized to surround an external
surface of said heart; said material having an elasticity and a
compliance reflecting a tendency to return to a rest state and an
ability to deform under strain, respectively; said elasticity and
compliance selected for said material to be placed surrounding said
external surface with a tightness selected for energy to be stored
in said material to assist chronic remodeling of said heart and to
avoid significant acute resistance to diastolic filling of said
heart.
2. The device according to claim 1 wherein said remodeling includes
a reduction in a volume of said heart.
3. The device according to claim 1 wherein said remodeling includes
a specific or directed altering of a shape of said heart.
4. The device according to claim 1 wherein the material has a
compliance greater than a compliance of a normal myocardium where
compliance is the inverse of stiffness, wherein said biocompatible
material conforms to an external surface of the heart and is sized
to provide resistance to circumferential expansion of the heart
without impeding systolic contraction.
5. The device according to claim 4, wherein the material has a
stiffness of less than about 3.8 lbs/in when subjected to a
uniaxial load at a strain of less than 30%.
6. The device according to claim 1 wherein the device has a
compliance greater than a compliance of a normal latissimus dorsi
where compliance is the inverse of stiffness, wherein said
biocompatible material conforms to an external surface of the heart
and is sized to provide resistance to circumferential expansion of
the heart without impeding systolic contraction.
7. The device according to claim 6, wherein the material has a
stiffness of less than about 0.5 lbs/in when subjected to a
uniaxial load at a strain of less than 30%.
8. The device according to claim 1, wherein the biocompatible
material has a stiffness of less than 0.5 lbs/in when subjected to
a uniaxial load at a strain of less than 30%.
9. The device according to claim 1, wherein the biocompatible
material has a stiffness between about 0.05 lbs/in and about 0.2
lbs/in when subjected to a uniaxial load at a strain of less than
30%.
10. The device according to claim 1, wherein the material has a
compliance that allows at least a 3% increase in volume for every 1
mm Hg change in applied pressure.
11. The device according to claim 1, wherein the material has a
compliance between about 5%/mm Hg and about 15%/mm Hg.
12. The device according to claim 1, wherein the material is sized
to be smaller than the external surface of the heart to which it is
applied, wherein the material is configured to exert a pressure on
the external surface of the heart that is no greater than an end
diastolic pressure of aright ventricle of the heart.
13. The device according to claim 12, wherein the material defines
a resting volume and is stretched to provide a stretched volume
that is at least 20% greater than the resting volume.
14. The device according to claim 1, wherein the material is sized
to be larger than the external surface of the heart to which it is
applied and adapted to be sized by adjustment during
implantation.
15. The device according to claim 1, wherein the material is
capable of an elastic recovery of at least about 50%.
16. The device according to claim 1, wherein the material is
capable of an elastic recovery of at least about 70%.
17. The device according to claim 1, wherein said material is
configured to cover a ventricular portion of the heart.
18. The device according to claim 1, wherein said material is
configured as a jacket having an upper and a lower end, wherein
said jacket is open at said upper and defines an internal volume
between said upper and lower end capable of receiving the
heart.
19. The device according to claim 18, wherein said material is
closed at said lower end.
20. The device according to claim 18, wherein said material is open
at said lower end.
21. The device according to claim 1, wherein said material is
configured to cover at least one ventricle of said heart.
22. The device according to claim 21, wherein said material is
configured to cover a left ventricle of said heart.
23. The device according to claim 21, wherein said material is
configured to cover a right ventricle of said heart.
24. The device according to claim 21, wherein said material is
configured to cover a right and a left ventricle of said heart.
25. The device according to claim 1, wherein said material is
configured as a patch.
26. The device according to claim 1, wherein said material
comprises intertwined fibers.
27. The device according to claim 26, wherein said material
comprises a knit.
28. The device according to claim 26, wherein said material
comprises a weave.
29. The device according to claim 1, wherein said material
comprises polyethylene terephthalate (PET) or other polymeric or
biological materials with said properties.
30. The device according to claim 26, wherein said fibers are
crimped.
31. The device according to claim 30, wherein said crimped fibers
are produced by stretch texturizing.
32. The device according to claim 31 wherein the fibers consist of
a stretch textured polyester or PET or other polymeric
materials.
33. The device according to claim 1, wherein said material is
configured to apply a pressure to the external surface of the heart
at end diastole of less than 10 mm Hg.
34. The device according to claim 1, wherein said material is
configured to apply a pressure to the external surface of the heart
at end diastole of between about 5 mm Hg and about 7 mm Hg.
35. The device according to claim 1, wherein the material is
applied over a pericardium of the heart.
36. The device according to claim 1 further comprising a less
compliant material sized to limit diastolic expansion beyond a
maximum.
37. A method for treating diseases of a heart, said method
comprising: surgically accessing the heart; applying a cardiac
support device to an external surface of the heart where said
cardiac support device includes: a biocompatible material sized to
surround an external surface of said heart; said material having an
elasticity and a compliance reflecting a tendency to return to a
rest state and an ability to deform under strain, respectively;
said elasticity and compliance selected for said material to be
placed surrounding said external surface with a tightness selected
for energy to be stored in said material to assist chronic
remodeling of said heart and to avoid significant acute resistance
to diastolic filling of said heart.
38. The method according to claim 37 wherein said remodeling
includes a reduction in a volume of said heart.
39. The method according to claim 37 wherein said remodeling
includes an altering of a shape of said heart.
40. The method according to claim 37 wherein said material has a
compliance greater than the compliance of a normal latissimus
dorsi, wherein said biocompatible material conforms to an external
surface of the heart and is sized to provide resistance to
circumferential expansion of the heart without impeding systolic
contraction.
41. The method according to claim 37, wherein the material is
configured to resist dilatation of the heart.
42. The method according to claim 37, wherein the material is sized
to be larger than the external surface of the heart to which it is
applied during diastole, said method further comprising a step of
adjusting a size of said material to conform to the external
surface of the heart.
43. The method according to claim 42, wherein the material is
adjusted to exert a pressure on the external surface of the heart
of less than about 10 mm Hg.
44. The method according to claim 42, wherein the material is
adjusted to exert a pressure on the external surface of the heart
of less than about 5 mm Hg.
45. The method according to claim 42, wherein the material is
adjusted to exert a pressure on the external surface of the heart
of less than about 2 mm Hg.
46. The method according to claim 37, wherein said material is
configured to acutely support the external surface of the heart
without custom fitting.
47. The method according to claim 46, wherein said material is
sized to be smaller than the external surface of the heart to which
it is applied during diastole.
48. The method according to claim 47, wherein said material exerts
a pressure on the external surface of the heart that is less than
an end diastolic pressure of a right ventricle of the heart.
49. The method according to claim 47, wherein said material is
configured to impose between about a 5% to about a 10% reduction in
maximum diastolic volume of the heart upon implantation.
50. The method according to claim 47, wherein said material is
configured to exert a pressure on the external surface of the heart
at end diastole between about 2 mm Hg and about 20 mm Hg.
51. The method according to claim 47, wherein said material is
configured to exert a pressure on the external surface of the heart
at end diastole between about 5 mm Hg and about 15 mm Hg.
52. The method according to claim 47, wherein said material is
configured to exert a pressure on the external surface of the heart
at end diastole between about 5 mm Hg and about 10 mm Hg.
53. The method according to claim 37, wherein said material is
configured to enhance a reduction in a size of the heart.
54. The method according to claim 53, wherein said material is
sized to be smaller than the external surface of the heart to which
it is applied during diastole.
55. The method according to claim 54, wherein the material defines
a resting volume, said method further comprising a step of
stretching the material to provide a stretched volume that is at
least 20% greater than the resting volume.
56. The method according to claim 55, wherein said material exerts
a pressure on the external surface of the heart that is less than
an end diastolic pressure of a right ventricle of the heart.
57. The method according to claim 54, wherein said step of
stretching the material provides the material with potential energy
to mechanically reduce a size of the heart.
Description
BACKGROUND OF THE INVENTION
[0001] Congestive heart disease is a progressive and debilitating
illness characterized by a progressive enlargement of the heart. As
the heart enlarges, the heart must perform an increasing amount of
work to pump blood each heartbeat needed for metabolism. In time,
the heart becomes so enlarged the heart cannot adequately supply
blood. An afflicted patient is fatigued, unable to perform even
simple exerting tasks and experiences pain and discomfort. Further,
as the heart enlarges, the internal heart valves may not adequately
close. This may impair the function of the valves and further
reduce the heart's ability to supply blood. Importantly, there is
no cure for congestive heart disease.
[0002] Congestive heart failure has an enormous societal impact. In
the United States alone, about five million people suffer from the
disease. Alarmingly, congestive heart failure is one of the most
rapidly accelerating diseases (about 400,000 new patients in the
United States each year). Economic costs of the disease have been
estimated at $38 billion annually. Not surprising, substantial
effort has been made to find treatments for congestive heart
disease.
[0003] Cardiomyoplasty is one potential treatment for moderate
stage congestive heart disease. In this procedure, the latissimus
dorsi muscle (taken from the patient's shoulder) is wrapped around
the heart and chronically paced to assist contraction of the heart
during systole.
[0004] One study speculates that an artificial elastic "sock" could
be used in place of the latissimus dorsi in adynamic
cardiomyoplasty. Kass et al., "Reverse Remodeling from
Cardiomyoplasty in Human Heart Failure," Circulation 91:9, May 1,
1995. Another study demonstrates that Bard Marlex sheets can be
used to wrap the heart as a substitute to the latissimus dorsi in
adynamic cardiomyoplasty. Oh et al., "The Effects of Prosthetic
Cardiac Binding and Adynamic Cardiomyoplasty in a Model of Dilated
Cardiomyoplasty," Journal of Thoracic Cardiovascular Surgery,
116:1, July 1998. German Utility Model Patent Application DE 295 17
393 U1 describes a pericardium prosthesis made from a
biocompatible, non-expansible material, or at least hardly
expansible material that surrounds the heart to prevent
over-expansion of the heart wall. PCT application WO 98/58598
describes an elastic pouch for at least partially enveloping a
heart. Commonly assigned U.S. Pat. No. 5,702,343 to Alferness dated
Dec. 30, 1997 teaches a jacket to constrain cardiac expansion
during diastole. Other teachings include those of commonly assigned
U.S. Pat. No. 6,123,662 and those of U.S. Pat. Application
Publication No. US 2002/0019580.
SUMMARY OF THE INVENTION
[0005] The invention provides a device for treating cardiac
disease. According to the invention, a highly compliant and elastic
device, constructed from a biocompatible material is applied to an
external surface of a heart. The device can be used to resist
dilatation of the heart, provide acute wall support, and/or to
enhance reduction in the size of the heart using stored potential
energy, without interfering with systolic contraction. According to
the invention, the device has a compliance that is greater than a
compliance of a normal myocardium, more preferably, the device has
a compliance greater than a compliance of a normal latissimus
dorsi. Considering that stiffness is the inverse of compliance, the
uniaxial stiffness of the material is generally less than about 0.5
lbs/in (i.e. load per width of device) when subject to a uniaxial
load at a strain of less than 30%, more typically between about
0.05 lbs/in and about 0.2 lbs/in. An alternative way to examine
compliance for a device that is applied to an enclosed volume is
based on a 3-dimensional volumetric compliance. The 3-dimensional
compliance of the device typically allows at least a 3% increase in
volume for every 1 mm Hg change in applied device pressure. More
typically, the material of the device has a 3-dimensional
volumetric compliance between about 5%/mm Hg and about 15%/mm Hg.
The device typically has an elastic recovery of at least about 50%,
but 70% to 100% is preferable.
[0006] In one embodiment, the material of the device is sized to be
smaller than the external surface of the heart to which it is
applied. In another embodiment, the material of the device is sized
to be larger than the external surface of the heart to which it is
applied and adapted to be sized by adjustment during implantation.
The material can be configured as a jacket for covering both
ventricles, one ventricle, ventricles and atria, atria or as a
patch covering a portion of a chamber.
[0007] The invention also provides a method for treating a cardiac
disease. The method includes a step of surgically accessing the
heart; and applying a cardiac support device to an external surface
of the heart. The method can be used to resist dilatation of the
heart, acutely support the wall of the heart, and/or to enhance
reduction in a size of the heart using stored potential energy.
BRIEF DESCRIPTION OF THE DRAWINGS
[0008] FIG. 1 is a schematic cross-sectional view of a normal,
healthy human heart shown during systole;
[0009] FIG. 1A is the view of FIG. 1 showing the heart during
diastole;
[0010] FIG. 1B is a view of a left ventricles of a healthy heart as
viewed from a septum and showing a mitral valve;
[0011] FIG. 2 is a schematic cross-sectional view of a diseased
human heart shown during systole;
[0012] FIG. 2A is the view of FIG. 2 showing the heart during
diastole;
[0013] FIG. 2B is the view of FIG. 1B showing a diseased heart;
[0014] FIG. 3 is a schematic showing the theory of operation of a
cardiac support device.
[0015] FIG. 4 is a perspective view of an embodiment of a cardiac
support device according to the present invention;
[0016] FIG. 4A is a side elevation view of a diseased heart in
diastole with the device of FIG. 4 in place;
[0017] FIG. 5 is a perspective view of another embodiment of a
cardiac support device with the apex open according to the present
invention;
[0018] FIG. 5A is a side elevation view of a diseased heart in
diastole with the device of FIG. 5 in place;
[0019] FIG. 6 is a plan view of an alternate embodiment of a
cardiac support device;
[0020] FIG. 7 is a side elevation view of a diseased heart in
diastole with another embodiment of a cardiac support device in
place;
[0021] FIG. 8 is a cross-sectional view of a device of the present
invention overlying a myocardium and with the material of the
device gathered for a snug fit;
[0022] FIG. 9 is an enlarged simplified view of a the fabric of a
knit construction at rest, suitable for use in the device of this
invention;
[0023] FIG. 10 shows compliance curves (pressure versus % volume
change) for a lower compliance (A) and a higher compliance (B)
device;
[0024] FIG. 11 shows compliance curves (pressure versus % volume
change) for a lower compliance device (A) and a higher compliance
device (B) at implant;
[0025] FIG. 12 is a device elastic potential energy plot comparing
the work energy stored at implant for a less compliant and elastic
device (A) and a highly compliant and elastic device (B);
[0026] FIG. 13 is a plot of the change in left ventricle end
diastolic dimension (LVEDD) over time from clinical studies;
[0027] FIG. 14 is a plot of device loading (.diamond-solid.) and
unloading (.box-solid.) for determining elastic recovery; and
[0028] FIG. 15 is a schematic of the mechanical roles in cardiac
support device therapy;
[0029] FIG. 16 is a plot of LVEDV (left ventricular end diastolic
volume) change results from a pre-clinical animal study with a
horizontal axis depicting time and with a vertical axis depicting
change in LVEDV from pre-implant conditions.
DESCRIPTION OF THE PREFERRED EMBODIMENT
[0030] A. Congestive Heart Disease
[0031] With initial reference to FIGS. 1 and 1A, a normal, healthy
human heart H' is schematically shown in cross-section and will now
be described in order to facilitate an understanding of the present
invention. In FIG. 1, the heart H' is shown during systole (i.e.,
high left ventricular pressure during the ejection phase). In FIG.
1A, the heart H' is shown during diastole (i.e., low left
ventricular pressure during the relaxation phase).
[0032] The heart H' is a muscle having an outer wall or myocardium
MYO' and an internal wall or septum S'. The myocardium MYO', septum
S' and valve plane VP' define four internal heart chambers,
including a right atrium RA', a left atrium LA', a right ventricle
RV' and a left ventricle LV'. The heart H' has a length measured
along a longitudinal axis AA'-BB' from an upper end or base B' to a
lower end or apex A'.
[0033] The heart H' can be visualized as having an upper portion
UP' and a lower portion LP', separated by the valve plane VP'. On
the external surface of the heart, the upper portion UP' and lower
portion LP' meet at a circumferential groove commonly referred to
as the A-V groove AVG' right and left atria RA', LA' reside in an
upper portion UP' of the heart H' adjacent the base B'. The right
and left ventricles RV', LV' reside in a lower portion LP' of the
heart H' adjacent the apex A'. The ventricles RV', LV' terminate at
ventricular lower extremities LE' adjacent the apex A' and spaced
therefrom by the thickness of the myocardium MYO'.
[0034] Extending away from the upper portion UP' are a plurality of
major blood vessels communicating with the chambers RA', RV', LA',
LV'. For ease of illustration, only the superior vena cava SVC' and
a left pulmonary vein LPV' are shown as being representative.
[0035] The heart H' contains valves to regulate blood flow between
the chambers RA', RV', LA', LV' and between the chambers and the
major vessels (e.g., the superior vena cava SVC' and a left
pulmonary vein LPV'). For ease of illustration, not all of such
valves are shown. Instead, only the tricuspid valve TV' between the
right atrium RA' and right ventricle RV' and the mitral valve MV'
between the left atrium LA' and left ventricle LV' are shown as
being representative.
[0036] The valves are secured, in part, to the myocardium MYO' in a
region of the A-V groove AVG' and referred to as the valve plane
VP' or valvular annulus VA'. The valves TV' and MV' open and close
through the beating cycle of the heart H'.
[0037] FIGS. 1 and 1A show a normal, healthy heart H' during
systole and diastole, respectively. During systole (FIG. 1), the
myocardium MYO' is contracting and the heart assumes a shape
including a generally conical lower portion LP'. During diastole
(FIG. 1A), the heart H' is expanding and the conical shape of the
lower portion LP' bulges radially outwardly (relative to axis
AA'-BB').
[0038] The motion of the heart H' and the variation in the shape of
the heart H' during contraction and expansion is complex. The
amount of motion varies considerably throughout the heart H',
although the external dimension of the heart H' generally reduces
from about 4% to about 10% from end diastole to end systole. The
motion includes a component which is parallel to the axis AA'-BB'
(conveniently referred to as longitudinal expansion or
contraction). The motion also includes a component perpendicular to
the axis AA'-BB' (conveniently referred to as circumferential
expansion or contraction).
[0039] Having described a healthy heart H' during systole (FIG. 1)
and diastole (FIG. 1A), comparison can now be made with a heart
deformed by congestive heart disease. Such a heart H is shown in
systole in FIG. 2 and in diastole in FIG. 2A. All elements of
diseased heart H are labeled identically with similar elements of
healthy heart H' except only for the omission of the apostrophe in
order to distinguish diseased heart H from healthy heart H'.
[0040] Comparing FIGS. 1 and 2 (showing hearts H' and H during
systole), the lower portion LP of the diseased heart H has lost the
tapered conical shape of the lower portion LP' of the healthy heart
H'. Instead, the lower portion LP of the diseased heart H bulges
outwardly between the apex A and the A-V groove AVG. So deformed,
the diseased heart H during systole (FIG. 2) resembles the healthy
heart H' during diastole (FIG. 1A). During diastole (FIG. 2A), the
deformation is even more extreme.
[0041] As a diseased heart H enlarges from the representation of
FIGS. 1 and 1A to that of FIGS. 2 and 2A, the heart H becomes a
progressively inefficient pump. Therefore, the heart H requires
more energy to pump the same amount of blood. Continued progression
of the disease results in the heart H being unable to supply
adequate blood to the patient's body and the patient becomes
symptomatic insufficiency. In contrast to a healthy heart H', the
external dimension of the diseased heart H generally reduces from
about 4% to about 6% from end diastole to end systole.
[0042] For ease of illustration, the progression of congestive
heart disease has been illustrated and described with reference to
a progressive enlargement of the lower portion LP of the heart H.
While such enlargement of the lower portion LP is most common and
troublesome, enlargement of the upper portion UP may also
occur.
[0043] In addition to cardiac insufficiency, the enlargement of the
heart H can lead to valvular disorders. As the circumference of the
valvular annulus VA increases, the leaflets of the valves TV and MV
may spread apart. After a certain amount of enlargement, the
spreading may be so severe the leaflets cannot completely close (as
illustrated by the mitral valve MV in FIG. 2A). Incomplete closure
results in valvular regurgitation contributing to an additional
degradation in cardiac performance. While circumferential
enlargement of the valvular annulus VA may contribute to valvular
dysfunction as described, the separation of the valve leaflets is
most commonly attributed to deformation of the geometry of the
heart H. This is best described with reference to FIGS. 1B and
2B.
[0044] FIGS. 1B and 2B show a healthy and diseased heart,
respectively, left ventricle LV', LV during systole as viewed from
the septum (not shown in FIGS. 1B and 2B). In a healthy heart H',
the leaflets MVL' of the mitral valve MV' are urged closed by left
ventricular pressure. The papillary muscles PM', PM are connected
to the heart wall MYO', MYO, near the lower ventricular extremities
LE', LE. The papillary muscles PM', PM pull on the leaflets MVL',
MVL via connecting chordae tendineae CT', CT. Pull of the leaflets
by the papillary muscles functions to prevent valve leakage in the
normal heart by holding the valve leaflets in a closed position
during systole. In the significantly diseased heart H, the leaflets
of the mitral valve may not close sufficiently to prevent
regurgitation of blood from the ventricle LV to the atrium during
systole.
[0045] As shown in FIG. 1B, the geometry of the healthy heart H' is
such that the myocardium MYO', papillary muscles PM' and chordae
tendineae CT' cooperate to permit the mitral valve MV' to fully
close. However, when the myocardium MYO bulges outwardly in the
diseased heart H (FIG. 2B), the bulging results in displacement of
the papillary muscles PM. This displacement acts to pull the
leaflets MVL to a displaced position such that the mitral valve
cannot fully close.
[0046] B. Cardiac Support Therapy
[0047] In general, cardiac support therapy uses a "passive"
mechanical implant to support the heart and resist circumferential
expansion of the heart during diastole and without actively
assisting contraction during systole. Herein, the term "passive" is
used to contrast the device with an "active assist" device which
uses supplied energy in order to operate, such as devices that
assist the heart in pumping blood flow into the aorta, for example,
left ventricular assist devices ("LVAD") and total artificial
hearts ("TAH"). However, the device of the invention does have some
mechanical components that involve energy input into the system,
and therefore are not entirely "passive." As used herein, the term
"active" refers to a device wherein energy is added to the system
on an ongoing basis. In contrast, a "passive" device, as used
herein, may use stored or potential energy. The potential energy
stored in the device is generally attributable to energy that is
input when the device is fit on the heart, which in part is due to
mechanical properties of the device (such as compliance and
elasticity). The device can be thought of as having stored energy,
similar to a pre-loaded spring. However, in contrast to an "active"
device, once the device of the invention is implanted, no
additional energy is added continually. The mechanical components
of the device that involve energy input are described in detail
below.
[0048] It is believed that the cardiac support device stimulates a
physiological response due to a mechanical effect, a
tissue-material interaction, or some combination thereof. While the
physiological response can be difficult to predict, the mechanical
interactions are more straightforward. FIG. 3 is a schematic
showing how a cardiac support device interrupts the cycle of heart
failure by disrupting excessive ventricular dilatation (i.e.,
abnormal dilation) during diastolic filling. Briefly, following an
injury to the myocardium, the heart's function may be reduced (A).
This stimulates a compensatory response of ventricular dilatation
(B) to improve output. However, ventricular dilatation causes
increased wall stress and stretch (D), which then triggers
neurohormonal activation (C), leading to modified gene expression
(E) which in turn leads to structural and functional changes in the
myocardium. These changes are also referred to as ventricular
remodeling (F). This further reduces cardiac function causing the
cycle to repeat with additional compensatory dilatation. FIG. 3
illustrates the potential benefits of a cardiac support device
providing wall support and resistance to ventricular dilatation. A
cardiac support device reduces the myocardial wall stress and
stretch (H), which helps to break the heart failure cycle and leads
to improved efficiency (I), reverse remodeling (J) and ultimately
improved cardiac function (K).
[0049] The mechanical effects that help interrupt the ongoing
ventricular dilatation in the heart failure cycle can be divided
into at least three mechanisms or components: (1) dilatation
constraint (acute and/or chronic); (2) acute wall support; and (3)
chronic potential energy release. As used herein, "dilatation
constraint" means resisting expansion or dilatation of a heart that
would result in a damaging increase in the volume of the heart. As
used herein, "acute wall support" means reducing stress on the wall
of the heart or supporting the internal pressure (i.e. reducing
transmural wall pressure) of the heart by off-loading the heart
acutely, at the time of the device implant. As used herein,
"chronic potential energy release" refers to the potential energy
of the device that is available (and released) to encourage
reduction in the size of the heart, in terms of volume and/or
dimension over time following implant of the device. Many of the
device parameters, whether used for dilatation constraint, acute
wall support and/or potential energy release, overlap. However, for
the sake of clarity the design characteristics and/or the method
for implanting a device for each of the three mechanisms or
components will be discussed separately below. FIG. 15 is a
schematic diagram showing the mechanical modes for the device
mechanisms and how they tie into the biological responses to
support the theory of operation in FIG. 3.
[0050] 1. Dilatation Constraint (Acute and/or Chronic)
[0051] Dilatation constraint refers to the resistance the device
provides to short-term transient dilatation and/or chronic cardiac
dilatation, in particular excessive ventricular dilatation.
Generally, dilatation constraint does not require energy input into
the device. The device is typically adjusted to conform to the
epicardial surface, to resist further dilatation of the heart.
"Acute" dilatation constraint refers to resistance to cardiac
dilatation from short term loading such as exercise loading.
Exercise loading refers to the loading that occurs when a person
performs a physical activity such as exercise. In exercise loading,
the heart increases it's volume to provide more output using the
Frank-Starling relationship. However, the increased volume results
in increased end diastolic loading of the ventricular wall due to
the Law of LaPlace. The Law of LaPlace is based on the concept that
the larger the vessel radius, the larger the wall tension required
to withstand a given internal fluid pressure. Larger ventricular
chamber volumes generally correspond to larger chamber radii.
"Chronic" dilatation constraint refers to resistance to continued
dilatation due to prolonged volume loading and cardiac remodeling.
Increased volume loading can also result from the intake of fluids,
which is not discussed here, or kidney damage that is also
associated with heart disease.
[0052] During dilatation constraint, the device reduces the
ventricular wall stress and stretch increase that accompany acute
and continued dilatation in heart failure. Both the reduction in
ventricular wall stress and stretch increase are "myocardial
displacement dependent." As used herein, the phrase "myocardial
displacement dependent" means that the amount of support or loading
provided by the cardiac support device is dependent on the amount
of myocardial wall dimensional dilatation caused by disease
progression or excessive loading. For this mechanical component,
the compliance of the device can be an important characteristic.
Generally, the compliance of the device can be important for acute
loading before the fibrosis encapsulates the device and for
long-term chronic dilatation. However, device compliance tends to
be less important for acute loading after the fibrosis develops.
Generally, lower compliance (i.e. higher stiffness) tends to
provide more resistance and support for dilatation.
[0053] Generally, support devices such as those mentioned in the
Background section of this application. have focused on the
mechanism of dilatation constraint.
[0054] Generally, when the device is used for dilatation
constraint, the device 10 surrounds the myocardium MYO, as shown,
for example, in FIG. 4. As used herein, "surround" means that the
device provides reduced expansion of the heart wall during diastole
by applying constraining surfaces at least at diametrically
opposing aspects of the heart. Generally, the diametrically opposed
surfaces are interconnected, for example, by a continuous material
that can substantially encircle the external surface of the
heart.
[0055] In one embodiment, the device is configured as a jacket 10
that defines a volume 16. Preferably the volume 16 is substantially
the same size as or larger than the volume of the heart H, in
particular the lower portion LP of the heart, at the completion of
systolic contraction such that the jacket 10 exerts no or only a
slight pressure on the heart at end systole. Generally, the jacket
10 is adjusted such that the jacket 10 resists enlargement of the
heart H during diastole without significantly assisting contraction
during systole. At time of placement, the device preferably exerts
no or only a small pressure on the heart H at end diastole of less
than 10 mm Hg, more preferably less than or equal to 5 mm Hg, most
preferably less than or equal to 2 mm Hg. Such pressure may be
determined by comparing load to the right ventricular end diastolic
pressure.
[0056] To permit the jacket 10 to be easily placed on the heart H,
the volume and shape of the jacket 10 may be larger than the lower
portion LP during diastole. So sized, the jacket 10 may be easily
slipped around the heart H. Once placed, the jacket's volume and
shape can be adjusted for the jacket 10 to snugly conform to the
external geometry of the heart H during diastole. For example,
excess material of the jacket 10 can be gathered and sutured S"
(FIG. 8) to reduce the volume of the jacket 10 and conform the
jacket 10 to the shape of the heart H during diastole. This shape
represents an adjusted volume. The jacket 10 resists enlargement of
the heart H beyond the adjusted volume without interfering with
contraction of the heart H during systole. As an alternative to the
gathering shown in FIG. 8, the jacket 10 can be provided with other
ways of adjusting volume. For example, as disclosed in U.S. Pat.
No. 5,702,343, the jacket can be provided with a slot. The edges of
the slot can be drawn together to reduce the volume of the jacket.
The volume of the jacket can be adjusted prior to, during, or after
application of the device to the heart.
[0057] Although, for dilatation constraint, the device is generally
adjusted to a snug fit as described above, it is also possible to
obtain the benefits of dilatation constraint using a device that
defines a volume that is smaller than the volume of the portion of
the heart H on which it is to be placed at end diastole. In this
embodiment, the device is stretched in order to place it around the
heart H, such that the compliance of the jacket 10 material and the
amount of expansion of the material at end diastole determine the
fit of the device without any further adjustment.
[0058] 2. Acute Wall Support
[0059] Acute wall support refers to a more immediate effect of the
cardiac support device. Generally, acute wall support is obtained
by adjusting the device such that the device applies an external
pressure to the heart. If desired, the device can be adjusted to
provide a dimensional reduction in the heart size. For example, the
device may be adjusted to slightly reduce cardiac dimension at the
time of implantation, preferably, no more than 10% reduction in
internal Left Ventricular End Diastolic Dimension (LVEDD). Thus,
rather than just reducing the increase in wall stress and stretch
due to dilatation constraint, energy is actively input at the time
of implantation to reduce the load on the wall acutely. However,
after the device is placed on the heart, no further external energy
is added. Thus, the device is still considered a "passive" device.
Acute wall support reduces wall stress (load dependent) and reduces
wall stretch (myocardial displacement dependent). As used herein,
the phrase "load dependent" means that the reduction in wall stress
is dependent on the amount of load applied, independent of the
amount of dimensional change. The reduction in end diastolic wall
stress is based on the change in transmural heart wall pressure. In
contrast, the amount of reduced wall stretch is related to the
dimensional reduction in the heart size. In contrast to a
dilatation constraint mechanism, energy is input at the time of
implantation for acute wall support and the material compliance is
less important. However, as mentioned previously, these components
overlap such that the benefits from both dilatation constraint and
acute wall support can be realized from the same device.
[0060] If the device is configured as a jacket 10, it may be
desirable to have a volume and shape that is larger than the lower
portion LP during diastole so that the jacket 10 may be easily
slipped around the heart H and adjusted (as with dilatation
constraint). However, it may also be desirable to use a device with
a volume and shape that is smaller than the lower portion LP of the
heart H during diastole. In this embodiment, the compliance of the
jacket 10 and expansion at diastole determine the fit, without
additional adjustment. When selecting or adjusting the jacket 10
for acute wall support, care should be taken to avoid impairing
normal cardiac function. During diastole, the left ventricle LV
fills with blood. If the jacket 10 is too tight, the left ventricle
LV may not adequately expand and left ventricular filling pressure
may rise. Furthermore, if the device encloses both ventricles such
as in FIG. 4A, care should be taken when selecting or adjusting the
jacket 10, because the wall of the right ventricle RV tends to be
thinner than the wall of the left ventricle LV and the pressure in
the right ventricle RV tends to be lower than the pressure in the
left ventricle LV. Preferably the pressure exerted by the jacket 10
on the heart H at end diastole is not greater than the end
diastolic pressure of the right ventricle RV. If the pressure
exerted by the jacket 10 is greater than the pressure of the right
ventricle RV, expansion and/or filling of the right ventricle RV
may be compromised. However, for a device that is applied to only
one of the ventricular chambers such as the Left Ventricle LV as
shown in FIG. 7, the pressure exerted by the jacket 10 at end
diastole is preferably less than the end diastolic pressure of the
LV.
[0061] Generally a jacket 10 that imposes between about a 5% to
about a 10% reduction in LVEDD (left ventricle end diastolic
dimension) serves to reduce cardiac volume without compromising
cardiac function. Preferably, the jacket 10 exerts pressure at end
diastole between about 2 mm Hg and about 20 mm Hg, more preferably
between about 5 mm Hg and about 15 mm Hg, and most preferably
between about 5 mm Hg and about 10 mm Hg, depending on the internal
end diastolic pressures of the heart chambers. The jacket may be
designed with multiple sections with different compliances and
pressures for a specific heart chamber.
[0062] 3. Chronic Potential Energy Release
[0063] In addition to dilatation constraint and acute wall support,
the cardiac support device may also be able to use stored potential
energy to enhance heart size reduction over time. Based on the Law
of LaPlace, reduced heart size will reduce the myocardial wall load
for a given internal chamber pressure. The chronic potential energy
release mechanism and the device properties that enable the size
reduction are key aspects of this invention. Generally, the
potential energy of the device is due to the fabric being stretched
at the time of implantation. Typically, the device is selected
and/or adjusted (if necessary) to have a "resting" size and/or
volume that are smaller than that of the enlarged heart to which it
is applied. Preferably the "resting" size of the device is
approximately the same size as the heart in a healthy state or some
other desired target size. As used herein the term "resting" means
that the fibers of the fabric are in a relaxed state such that
energy is not required to keep the fibers in the "resting" or
"relaxed" state. When the material is "stretched" to accommodate
the enlarged heart, work energy must be input to create the
"stretched" configuration. The amount of energy input and stored in
the device is based on the amount of strain (or stretch) and the
load required to obtain that strain. According to one aspect of the
invention, the material is stretched during implantation, wherein
the stretching provides the material with potential energy that can
be used to enhance reduction in the size of the heart. In one
embodiment, the material is stretched to provide a stretched volume
that is at least 20% greater than the resting volume, more
preferably, the material is stretched to provide a stretched volume
that is about 40% greater than the resting volume, more preferably
about 60% greater than the resting volume. The maximum stretch
should be based on the limit of heart volume reduction desired.
Similarly, a patch device that covers a small area (i.e. FIG. 6)
rather than encapsulating a volume may have similar stretch targets
based on area or length rather than volume.
[0064] Again, as with acute wall support, care should be taken to
avoid exerting too much pressure on the heart, such that cardiac
function is impaired. For this mechanical mechanism or component,
the device preferably exerts pressures similar to those described
for the acute wall support mechanism. Lower pressures may be
effective and more preferred depending on the compliance and
elasticity of the device and the desired level of stored potential
energy.
[0065] Both the compliance and elasticity of the material are
important parameters for the chronic potential energy release
mechanism. Compliance refers to the ability of the device to deform
under load. In engineering terms it is the inverse of stiffness.
Elasticity refers to the ability of the device to return to its
original dimension upon unloading after being deformed by a load.
The compliance of the device and the load applied determine the
amount of energy added to the system at the time of implant.
However, the device elasticity determines the new resting state of
the device after it has been stretched out for implantation, and
how much stored potential energy can be released from unloading the
device. Once the device is implanted, the heart will generally
reduce in size over time to reduce the external load applied by the
device. The amount of potential energy stored and the elasticity in
the device will affect how much the device can mechanically reduce
and reshape the heart from a dilated size to possibly a normal
size. A device having high compliance and high elasticity is
generally preferred for this mechanism to increase the amount of
potential energy that can be stored and recaptured.
[0066] It is noted that the tissue response to the implanted device
may cause the device to be encapsulated in a thin layer of
fibrosis. The collagenous fibrotic tissue can be remodeled when it
is subjected to chronic loads. Thus, after the device is
encapsulated by fibrosis, the composite compliance of the fibrosis
and device may be reduced for short-term transient loads. However,
for long-term loads such as reduction in the heart size due to the
chronic potential energy release of the device, fibrosis is
believed to have only a minor or insignificant impact on the
compliance and elasticity of the device. Over time the fibrosis is
unlikely able to support the load from the heart or the device.
Thus, the fibrosis tends to remodel as the heart reduces over time
and the compliance and elasticity of the device continue to drive
the mechanical reduction in heart size until the potential energy
of the device is fully released or the heart size stabilizes.
[0067] The chronic potential energy release mechanism of the device
reduces wall stress and reduces wall stretch, both of which are
myocardial displacement dependent. As the device mechanically
causes the heart to reduce in size, the heart wall stress and
stretch reduce due to the change in geometry. The chronic potential
energy mechanism was demonstrated in a pre-clinical animal model.
FIG. 16 with the early results of an animal study using canines
with failing hearts shows significantly larger left ventricular end
diastolic volume (LVEDV) reduction in two animals implanted with a
higher compliance device (A) when compared to six animals implanted
with the current lower compliant device (B). All animals were
implanted with similar loading and little to no acute reduction at
the time of implant. The additional potential energy stored in the
high compliance devices was able drive the size reductions by over
3 times more volume.
[0068] C. Cardiac Support Device.
[0069] The invention provides a device having a compliance and/or
elasticity to render it suitable for use for one or more of the
following treatments: resisting enlargement of the cardiac
dimension (dilatation constraint), offloading stress from the
myocardial wall (acute wall support), and enhancing reduction in
cardiac dimension (chronic potential energy release).
[0070] Generally, the device is configured to cover at least part
of the epicardial surface, typically at least the ventricles. As
used herein, the term "cover" means that the device is in contact
with an external surface and applies a force on the surface of the
heart. Generally, the device contacts an epicardial surface of the
heart, but it can also be applied over the pericardium.
[0071] A device that "covers" the lower extremities of the heart
may be constructed as a continuous material that can substantially
encircle, or "surround", the external surface of the lower
extremities of the heart (See, FIGS. 4, 4A, 5, 5A). In an alternate
embodiment, the device provides for localized support of the heart,
particularly during diastole. According to this embodiment, a
device 10 may be configured as a "patch" (See, FIG. 6). A patch may
be useful to provide dilatation constraint or acute wall support
over a localized area of injury such as an acute myocardial
infarction (AMI) or a wall aneurysm. In the case of an aneurysm, it
may be advantageous to take advantage of the chronic potential
energy release mechanism to restore the wall shape over time. When
discussing a "patch", the size of the patch is selected to cover an
area of the epicardial surface of the heart without completely
surrounding the circumference of the heart. In yet another
embodiment, the device may be configured to cover only a left or
right ventricle (See, FIG. 7). Typically, in this embodiment, the
device is attached to the heart proximate the septal wall S'. If
desired, the device can be constructed from material having one or
more compliances or be constructed as one or more separate
components. The mechanical characteristics of each component may be
designed to specifically target one or more of the mechanical
mechanisms of device therapy previously described. With reference
now to FIGS. 4, 4A, 5 and 5A, the device of the present invention
is shown as a jacket 10 of flexible, biologically compatible
material. As used herein, the term "biologically compatible
material" refers to material that is biologically inert such that
the material does not adversely result in excessive injurious
responses such as chronic inflammation which would adversely affect
the myocardium and potentially surrounding tissues.
[0072] A jacket 10 is an enclosed material having upper and lower
ends 12, 14. The jacket 10, 10' defines an internal volume 16, 16'
which is completely enclosed but for the open ends 12, 12' and 14'.
In the embodiment of FIG. 4, lower end 14 is closed. In the
embodiment of FIG. 5, lower end 14' is open. In both embodiments,
upper ends 12, 12' are open. Throughout this description, the
embodiment of FIG. 4 will be discussed. Elements in common between
the embodiments of FIGS. 4 and 5 are numbered identically with the
addition of an apostrophe to distinguish the second embodiment and
such elements need not be separately discussed.
[0073] The jacket 10 is dimensioned with respect to a heart H to be
treated. Specifically, the jacket 10 is sized for the heart H to be
enclosed within the volume 16. The jacket 10 can be slipped around
the heart H. The jacket 10 has a length L between the upper and
lower ends 12, 14 sufficient for the jacket 10 to enclose the lower
portion LP. In one embodiment, the upper end 12 of the jacket 10
extends at least to the valvular annulus VA and further extends to
the lower portion LP to enclose at least the lower ventricular
extremities LE. If desired, the jacket 10 may be sized so that the
upper end 12 resides in the A-V groove AVG. Where it is desired to
treat the upper portion UP, the jacket 10 may be extended to cover
the upper portion UP.
[0074] After the jacket 10 is positioned on the heart H as
described above, the jacket 10 is secured to the heart. Preferably,
the jacket 10 is secured to the heart H through sutures. The jacket
10 is sutured to the heart H at suture locations 15
circumferentially spaced along the upper end 12. While a surgeon
may elect to add additional suture locations to prevent shifting of
the jacket 10 after placement, the number of such locations 15 is
preferably limited so that the jacket 10 does not restrict
contraction of the heart H during systole. Other attachment methods
such as staples or clips may be acceptable as an alternative to
sutures along the upper end 12.
[0075] The jacket 10 can be adjusted to provide the appropriate fit
after placement around the heart. Alternatively, the jacket 10 can
be sized to obtain the appropriate fit based on device compliance,
desired level of fit and the size of the portion of the heart H the
device is intended to cover.
[0076] a. Compliance
[0077] As used herein, the term "compliance" refers to the load
required to deform the material. As mentioned previously, in the
field of engineering it is the inverse of stiffness. The compliance
can be described in terms of displacement (inches or centimeters),
strain (inch/inch or cm/cm) or volume (in.sup.3, cm.sup.3 or ml)
per a unit load (pounds or kilograms) or pressure (psi or mm Hg).
The compliance of the cardiac support device can have a significant
impact on the mechanical mechanism and effectiveness in the
therapy, as well as allowing it to be stretched for accommodating
an enlarged heart. It should also be noted that compliance is not
necessarily constant over a given range of displacement. In fact,
compliance that decreases with increased stretch is a common
characteristic of many materials.
[0078] Due to the Frank-Starling behavior of the heart, a cardiac
support device that has less compliance than the myocardium at
small deformations may not be desirable. Generally, the cardiac
output demand for the heart changes depending on physical activity.
To increase the cardiac output according to the Frank-Starling
mechanism, the pre-load or end diastolic volume of the heart is
increased such that the muscle fibers are temporarily stretched.
The stretching of the muscle fibers helps increase heart capacity
and myocyte contractility and therefore cardiac output. If a
cardiac support device with less compliance than the myocardium is
applied to the surface of the heart, the heart may not be able to
utilize the Frank-Starling mechanism effectively without increasing
ventricular filling pressure. Thus, ventricular filling may be
negatively impacted and mimic a cardiovascular disease known as
constrictive pathologies. Thus, a cardiac support device with a
higher compliance than the myocardium is generally preferred.
[0079] Other evidence indicating that cardiac support with higher
compliance may be preferable can be found in examining the
compliance of the normal pericardium and the latissimus dorsi
muscle used to wrap the heart for cardiomyoplasty. The stiffness of
living myocardium and latissimus dorsi muscle is complex and has
both active and passive elements. For simplicity, only the passive
elements will be examined.
[0080] Table 1 contains a comparison of passive stiffness of
myocardial tissue, pericardial tissue, latissimus dorsi muscle
tissue and a sample cardiac support device that has been described
in previous patents (i.e. U.S. Pat. No. 6,085,754 and International
patent application publication No. PCT WO 01/95830) (this sample
cardiac support device is referred to herein as the "prior knit
device"). The values in Table 1 were derived based on uniaxial
loads at strains less than 30%. As shown in Table 1, the
pericardial tissue is much more compliant than the myocardium for
low strains, but stiffens at higher strains to become less
compliant than the myocardium. If the pericardium was as stiff for
low strains as higher strains, ventricular filling would probably
be impaired, similar to constrictive pericarditis or cardiac
tamponade. The stiffness comparisons provided in Table 1 illustrate
that the latissimus dorsi (sometimes used to wrap the heart for
cardiomyoplasty) is also more compliant than the myocardial tissue.
The sample cardiac support device (i.e., the prior knit device)
also has a greater compliance than the myocardium and similar, but
slightly less compliance than the latissimus dorsi muscle. As
mentioned earlier in the Background section, Oh et al. used a very
non-expansible material to wrap around the heart known as Bard
Marlex. Although the Bard Marlex helped to limit progressive
dilatation in this study, it was not as effective as the latissimus
dorsi muscle in adynamic cardiomyoplasty. The uniaxial stiffness of
Bard Marlex has been measured to be less compliant than the
myocardium as shown in Table 1. The data in Table 1 thus supports
the concept that a myocardial support device should preferably be
more compliant than the myocardium.
1TABLE 1 Relative (uniaxial) Stiffness* Component (lbs/in)
Reference Myocardium 3.8 to 5.0 Sideman & Beyar, "Simulation
and Control of the Cardiac System," CRC Press, Inc., 1987, Chapter
5. Normal Human 0.1 to 25.0 Lee et al., "Biaxial mechanical
properties of human Pericardium of human pericardium and canine
comparisons, "Am. J. Physiol., 1987, 253:H75-H82. Latissimus Dorsi
0.5 to 0.7 Reichenbach et al., "Passive characteristics of
conditioned skeletal muscle for ventricular assistance," ASAIO J.,
1999 Jul.-Aug.;45(4):344-9. Cardiac Support 0.8 to 1.7 Bench
testing Device (i.e., the prior knit device) Bard Marlex 5.9 to
25.0 Bench testing *Notes: Slope of stress versus strain curve
(i.e..sigma./.epsilon.) is a measure of stiffness (lbs/in.sup.2).
Incorporating the material thickness (t), a measure of relative
stiffness is given by .sigma./.epsilon. (lbs/in). This is a measure
of load per inch width of material required to produce a given
strain. Compliance is the inverse of stiffness. The uniaxial
stiffness for various materials were derived from references listed
for strains up to 30%.
[0081] In the Kass et al. article mentioned in the Background
section, it was speculated that an artificial elastic "sock" could
be used to replace the latissimus dorsi muscle in adynamic
cardiomyoplasty. This reference seems to use the term elastic
relative to compliance (not in it's true engineering sense) and
makes the comparison to replacing the latissimus dorsi with a
device of similar compliance. The cardiac support device in Table I
has compliance that is comparable or slightly less than the
latissimus dorsi muscle. However, the inventors have found that a
high compliance cardiac support device may have superior
performance. The advantages of a high compliance cardiac support
device (i.e., a device having a compliance greater than the
latissimus dorsi) are not disclosed by Kass et al.
[0082] Compliance of cardiac support devices can be measured in
vitro to determine either uniaxial directional compliance or
3-dimensional full device volumetric compliance. The uniaxial
directional compliance can be determined by taking samples of a
selected device. These samples can be mounted on a standard
hydraulically actuated tensile testing machine such as those
supplied by MTS Systems Corporation or Instron Corporation. The
compliance or stiffness characteristics of the device are obtained
by measuring the load versus deflection of the sample. The device
and Bard Marlex stiffness provided in Table 1 were determined using
this method for comparison purposes.
[0083] In use, the compliance of the cardiac support device is more
realistically based on 3-dimensional loading than uniaxial loading.
Thus, an in vitro test was developed to examine full device
compliance. For this test, the circumference of the base end of a
sample cardiac support device (i.e., the prior knit device) was
mounted to a support plate to simulate the attachment of a device
near the heart valve plane. A balloon bladder was placed in the
volume defined by the device and filled with saline to simulate the
external heart ventricular volume. To offset the effect of the
weight of fluid within the balloon, the mounted device and balloon
are suspended in a tank of saline maintained at approximately
37.degree. C. The apex of the device is supported so when the fluid
is added to the balloon, the device expands primarily
circumferentially, to better simulate the dilatation of a heart in
failure. The internal volume of the balloon was monitored by
recording the volume of fluid that was added incrementally. At each
fluid increment the pressures within the balloon are monitored
using a Millar catheter tip transducer and between the device and
the external surface of the balloon using a "pillow" device and
methods similar to those described by Tyberg et al. ("Static and
dynamic operating characteristics of a pericardial balloon,"
Hamilton D R, Devries G, Tyberg J V, J Appl Physiol, April
2001;90(4):1481-1488). Both the internal balloon pressure and
pillow pressures track very closely, indicating very little
resistance from the balloon.
[0084] Typical compliance curves obtained using this method with
normalized percentage volume changes are shown in FIG. 10. FIG. 10
illustrates the 3-dimensional or volumetric compliance curves for
two cardiac support device configurations. Curve A illustrates a
lower compliance device, while curve B represents a higher
compliance device.
[0085] Data indicates that a high compliance cardiac support device
may be desirable in many circumstances. As discussed above, the
compliance of a cardiac support device can vary over a given range
of displacement, or depending whether or not the device is subject
to uniaxial or multiaxial loads.
[0086] As used herein, the term "high compliance cardiac support
device" refers to a device having a compliance that is greater than
that of a normal myocardium, and more preferably greater than the
compliance of the latissimus dorsi muscle. In one embodiment, or
characterization, the high compliance device of the invention can
thus be described as having a stiffness less than 3.8 lbs/in for
uniaxial strains up to 30%. As shown by the data in Table 1 and the
discussions above, it may be more preferable that the device has a
compliance that is greater than a normal Latissimus Dorsi muscle,
i.e., a stiffness less than 0.5 lbs/in for uniaxial strains up to
30%. Typically, when referring to a "high compliance" device
herein, the inventors are referring to a cardiac support device
having a stiffness less than about 0.5 lbs/in when subjected to a
uniaxial load at strains up to 30%, more typically between about
0.05 lbs/in and about 0.2 lbs/in. It will be appreciated that the
foregoing description of data for strain up to 30% is intended to
be representative and not to suggest strains greater than 30% are
not applicable to the present invention.
[0087] Another way in which the compliance of the high compliance
device can be characterized is based on 3-dimensional volumetric
compliance in terms of the percentage of volume increase (%) over
applied pressure (mm Hg). Using this characterization in a
representative example, the high compliance device of the invention
will have a compliance that allows at least a 3% increase in volume
for every 1 mm Hg increase in pressure (3%/mm Hg), more preferably
between about 5%/mm Hg and about 15%/mm Hg. Actual volumes will
depend upon the specific compliance. Again, it will be appreciated
the foregoing is a non-limiting example.
[0088] i. Material
[0089] The high compliance cardiac support device of the invention
can be fabricated using various materials and configurations to
provide the mechanical characteristics desired. In a preferred
configuration, the device is constructed from a warp knitted fabric
18 of polyester fibers. Generally, the fabric 18 material is formed
from intertwined fibers 20 that are made up of a plurality of
filaments 30, as shown in FIG. 9. The compliance of the material
may be due to a variety of factors, including, but not limited to,
the compliance of the individual filaments 30 that make up the
fibers 20 (see section b. Elasticity), the relative movement of the
filaments 30 within a fiber 20, and/or the relative movement of the
intertwined fibers 20 when subjected to load. Texturizing of the
yarn can impact the compliance and elasticity of the fibers.
Preferably, the fiber material and texturizing result in a
compliant and elastic fiber such as a stretch polyester.
[0090] Compliance due to the relative movement (e.g., geometric
deformation of the fabric openings) of the intertwined fibers 20
may be affected by the manner in which the fibers 20 are entwined.
For example, a knit material will tend to be more compliant than a
woven material because the loops of the knit are capable of
deforming (e.g., widening or lengthening) to accommodate applied
stress. In comparison, woven materials tend to have less elongation
unless elastomeric fibers are used. Knit material also tends to
recover well from deformation because the loops attempt to return
to their original positions. The looped configuration of the fibers
accommodates this recovery more readily than does the interwoven
configuration found in woven materials. The ease and quickness with
which elastic recovery takes place is also dependent on the fiber
composition. The fibers 20 of the jacket 10 material may be
entwined as a knit (for example, a warp knit) or as a weave.
Preferably, the fibers 20 of the jacket 10 material are entwined as
a knit.
[0091] ii. Manufacturing a High Compliance Device
[0092] The compliance of the cardiac support device can be due to
the intertwining of the fabric fibers, or due to the
compliance/elasticity of the fibers themselves, as discussed above.
Additionally, the compliance of the cardiac support device can be
altered by the method of processing the fabric.
[0093] For example, the compliance of the material of a cardiac
support device can be increased by "shrinking" the material of the
device, such that the device then includes more material within the
same unit area and the fibers are closer together and more
compressed, as compared to the device before the "shrinking"
process. For example, shrinking can be accomplished by heating the
device. A memory condition can be introduced by a high temperature
exposure or set temperature within the fibers, which modifies the
"resting" state of the fibers (i.e., the state to which they
naturally return without the use of force). Exposure to
temperatures below the set temperature can cause the fibers to
respond by shrinking to the at rest condition. However, exposure to
new temperature conditions above the original set temperature while
subjected to a load will create a new at rest configuration.
Additionally, changing the fabric knit configuration, fiber
texturizing or fiber material can further increase the compliance
of the original device.
[0094] Thus, in one embodiment, a high compliance device is
manufactured by adding additional material to a fabric pattern of a
lower compliance device (also referred to as the "original"
device). Both patterns are shrunk to the same size, for example,
using a heat set mandrel (i.e., the same heat set mandrel is used
for the "original device" and the "highly compliant" device). This
method can easily increase the compliance of the device 5 to 10
times (at low to moderate strains) over the original device.
[0095] In manufacturing, the device is shaped to that of a healthy
heart so that the device not only uses its stored energy to reduce
size but also to help the patient's heart restore shape. Both
beneficial attributes are referred to as remodeling.
[0096] b. Elasticity
[0097] As used herein, the term "elastic" refers to the ability of
the deformed material to return to its initial state after a
deforming load is removed. A device that is highly elastic can
undergo very large deformations, but upon unloading returns to or
close to its original state. With respect to a cardiac support
device, elasticity may be important to the cardiac support device
for maintaining an external load on the heart as it reduces in
size.
[0098] When a material is subjected to a deformation, the
deformation is either plastic or elastic. If the deformation is
plastic, it does not rebound when unloaded. The degree of
elasticity for a given loading can be characterized as the
percentage of the deformation that rebounds upon unloading. Thus
when unloaded, an entirely elastic material would rebound to its
original state and characterized as 100% elastic at that load.
Whereas, a material that does not rebound at all from its deformed
state would be considered to have undergone an entirely plastic
deformation and would be considered 0% elastic at that load. In
general, the amount of elastic recovery for the cardiac support
device (in %) can be calculated as 100%(d.sub.1-d.sub.2)/d.sub.1,
where d.sub.1 is the initial deformation and d.sub.2 is the
deformation after unloading. The deformations d.sub.1 and d.sub.2
can be based on any dimensional measure of length, area or volume
as long as the units are consistent.
[0099] Preferably the device 10 has an elastic recovery of at least
about 50 %. However, it should be at least enough to allow the
device to deform elastically up to the desired reduction in cardiac
dimension targeted for the chronic potential energy release. Thus,
if the device is implanted at 50% fabric strain and the desired
heart size is calculated to be at a point of 25% fabric strain, it
would be preferable to have a device capable of at least about 50%
elastic recovery, more preferably at least about 70% elastic
recovery.
[0100] As with compliance, the elasticity of the material may be
due to a variety of factors. The elasticity of the base material
used to fabricate the device is one factor in determining the
elastic recovery. For the cardiac support device, one suitable
material is polyethylene terephthallate (PET), more commonly known
as polyester. Other biologically compatible materials could also be
used to provide the desirable amount of elasticity. In addition to
the base material, the configuration and heat-induced memory are
also important in determining the elasticity of the device. In one
embodiment, a warp knitted fabric fabricated from continuous
multi-filament set textured yarns is used. The fabric knit
configuration contributes to the elastic performance of the device
as well as its compliance. However, the process of texturizing the
yarn fibers 20 introduces a permanent crimp in the yarn that is
very important to the compliance and elastic performance of the
final device.
[0101] The permanent crimp induced in the individual filaments 30
that make up the yarn fibers 20 during texturizing provides a
memory to the yarn. The permanent crimp can be deformed when
loaded, but will have a tendency to return to the crimped
configuration when unloaded (i.e. elastically recover). The
texturizing process generally involves heat and deformations to
form the permanent crimp. Stretching the fabric and heating to a
higher temperature during the final processing of the device can
remove some of the yarn crimp and provide a new memory
condition.
[0102] The preferred permanent yarn crimp for the original cardiac
support device is produced by set texturizing the yarns, then
processing the final device by heat setting it with the device
slightly stretched. Increased compliance and elasticity can be
obtained using the same polymer and fabric knit configuration, but
with no final device heat set or by using other texturizing
processes such as stretch textured yarns. As mentioned, in the
preferred configuration elastic recovery is at least 50%, but most
preferable 70% to 100%.
[0103] Device elasticity can be determined from the compliance
curves for loading and unloading a device. The in vitro 3D balloon
compliance test described in the previous section can be used to
load and unload the device to determine the elastic rebound.
[0104] D. Benefits
[0105] The device 10 of the invention may provide some or all of
the following benefits.
[0106] 1 Reduction in Heart Dimension
[0107] The device 10 of the invention is a highly compliant and
elastic device that is capable of mechanically reducing the heart
size over time by using the chronic potential energy release
mechanism previously described. The reduced heart size is
beneficial due to reduced wall stresses, which may, in turn, lead
to improved cardiac function. The benefit of reducing heart size
with a high compliance device can be illustrated by comparing a
lower compliant device to a high compliance device.
[0108] FIG. 11 shows the compliance curves of two devices as
implanted over the ventricular portion of the heart. The lower
compliance (A) and higher compliance (B) devices are both stretched
to apply the same external pressure (approximately 6 mm Hg) to the
ventricular portion of the heart at the time of implant and initial
heart volume (i.e. 0% heart volume increase). The devices plotted
in FIG. 11 are the same devices as shown in FIG. 10. The zero
points on the horizontal axes on the two Figures are not the same.
Therefore percent calculations between the Figures will differ. In
FIG. 10, the zero point is an "at rest" value for the device. In
FIG. 11, the zero point is after implantation. The higher
compliance of the "B" device was obtained by adding more material
to the device so that less yarn crimp was removed during heat
setting. The lower compliance of the "A" device is indicated by a
steeper line. The compliance of either device at any point of
either curve can be expressed as 1/slope of the curve at that
point. Thus, at implantation, the compliance of the higher
compliance device (B) is 5.5%/mm Hg compared to about 2%/mm Hg for
the less compliant (A) device. If the operating range of the device
is assumed to be below a 20% volume increase, the compliance range
is between about 3%/mm Hg and 20%/mm Hg for the highly compliant
(B) device and between about 1%/mm Hg and 3.5%/mm Hg for the less
compliant (A) device. In this example, the higher compliance device
is approximately 3 to 6 times more compliant depending on the given
condition within the operating range. As used herein, a "high
compliance" device refers to a device having a compliance between
about 3%/mm Hg and about 20%/mm Hg , or greater. A "low compliance"
device refers to a device having a compliance between about 1%/mm
Hg and about 3%/mm Hg, or lower. In FIG. 11, the devices can
theoretically apply an external pressure to the heart until the
heart volume decreases to the point where the compliance curve
crosses 0 mm Hg. These curves are based on loading, not unloading.
Therefore, as the heart volume decreases, these curves assume that
both devices have 100% elastic recovery. In general, even though
the two devices do not have 100% elastic recovery, for comparison
purposes the higher compliance device will have better elastic
recovery than the low compliance device. Whereas a "low compliance"
device may have the potential to reduce the size of a heart between
about 10% to about 20% in volume, a high compliance device can
continue to apply an external load to the heart to achieve up to
between about a 50% to about a 60% volume reduction.
[0109] Depending on the heart shape change that is assumed (i.e.,
cylindrical or spherical), the volume decrease for the "low
compliance" device corresponds to a decrease in diameter between
about 5% to about 10%. The volume decrease for the "high
compliance" device similarly corresponds to a decrease in diameter
between about 15% to about 30%.
[0110] A "low compliance device" corresponding to the lower
compliance device (A) has been implanted in human clinical trails
with follow-up out to 12 months post-implant (i.e., the prior knit
device). FIG. 13 shows the average change in left ventricular end
diastolic diameter (LVEDD) for 17 patients receiving the lower
compliance device. At implant the hearts were fit to provide acute
support that resulted in a 5.2% reduction in LVEDD. After 3 months
post-surgery, the LVEDD decrease another 4.8% on average. This
additional chronic reduction in LVEDD corresponds closely with the
5% to 10% diameter reduction of the external ventricular size
predicted by the device compliance curve shown in FIG. 11 and the
chronic potential energy mechanism. In fact, the amount of elastic
recovery for the low compliance device has been measured in vitro
to be approximately 70% to 80%, depending on the loads applied.
[0111] A typical loading and unloading curve for a lower compliance
device is shown in FIG. 14. The elastic recovery calculated from
FIG. 14 is approximately 80%. This was calculated based on the
percentage of the volume change from initial to fully deformed that
was recovered (i.e. fully deformed to new unloaded resting state
volume). If the predicted diameter reduction range of 5% to 8% is
reduced to account for less than 100% elastic recovery, the
expected decrease in diameter would be between 3.5% and 6.5%. The
actual clinical result is nearly in the middle of this predicted
range.
[0112] The potential reduction in heart size attributable from the
chronic potential energy release mechanism can also be examined
based on the energy that is stored in the device relative to the
device compliance. Elastic potential energy stored in a spring is
equal to the amount of work energy (U) used to compress it if no
frictional or other losses are assumed. Thus, the work energy can
be determined as follows:
U=work energy=Fx/2
[0113] Where:
[0114] F=applied force=Kx
[0115] K=stiffness=1/compliance
[0116] X=displacement
[0117] FIG. 12 illustrates the potential energy stored in both a
low compliance (A) and high compliance (B) device at implant. Both
energy curves assume that the device is implanted on the same size
heart (i.e. external diameter of 8.4 cm.) with the same externally
applied pressure of 6 mm Hg when implanted. The high compliant
device has nearly 4 times more energy (508 mJ versus 131 mJ) at the
time of implant. Although the amount of energy stored is due to the
compliance, the amount available for release to reduce the heart
size is based on the elastic recovery of the device. For example,
if the device has 80%,elastic recovery, then 80% of the energy is
available to drive the heart smaller, while 20% of the energy is
lost to permanent deformation of the material.
[0118] 2. Eliminate Surgical Fit
[0119] As discussed previously, in one embodiment, the cardiac
support device is adjusted at the time of implantation to provided
the desired fit. The adjustment of the device allows it to be used
on dilated hearts having a large range of shapes and sizes. To
accommodate such variability in dilated hearts, the device is
manufactured in many sizes. However, many dozens of sizes would be
necessary to provide a sufficient selection for providing the
appropriate fit across all the range of heart shapes and sizes.
[0120] One advantage of a "high compliance" jacket is that each
jacket can adapt to a large shape/size range, yet still provide the
appropriate fit. Since compliance is defined as the deformation for
a given load, a high compliance device will result in a large
change in deformation with a relatively small change in load. Thus,
a target load or fit range can theoretically be accommodated by a
larger displacement range with a high compliance device than for a
low compliance device.
[0121] For example, two device compliance curves, low compliance
(A) and high compliance (B) are shown in FIG. 10. If the chosen
target fit load (i.e. pressure applied to the epicardial surface of
the heart) is between about 5 mm Hg to about 7 mm Hg of pressure,
the dimensional stretch range from the device resting state can be
determined from the curves for each device. Device "A" can be
stretched anywhere between a 17% to 23% (6% range) increase in
volume from it's starting volume and will apply a 5 mm Hg to 7 mm
Hg pressure. However, the higher compliance device "B" can be
stretched over a larger range of 129% to 156% (27% range) from it's
starting volume for the same resulting load. Now suppose it is
desirable to manufacture devices that will apply a 5 mm Hg to 7 mm
Hg load for heart sizes from 645 ml to 780 ml of external
ventricular volume. The high compliance device (B) would require
only one size to cover the range of heart sizes selected, but it
would take 4 sizes of the low compliance device (A) to accommodate
the heart size range. This example is illustrated in Table 2.
2TABLE 2 No. Device Size (@ rest) Min. Volume Max. Volume Sizes
Device (ml) (ml) (ml) Required B 500 645 780 1 A 550 644 677 4 578
676 711 607 710 746 637 745 783
[0122] Eliminating surgical fit based on a high compliance device
may be beneficial for several reasons. First of all, although the
surgical procedure for implanting a low compliance device is
relatively simple compared to other cardiac surgeries, eliminating
the fitting process would further simplify the surgery. The
surgical fitting step is one of the most time-consuming steps of
the surgical implant process. Eliminating this step could shorten
the overall surgical time. This would result in the patient
undergoing anesthesia for a shorter period of time, reducing the
risks due to anesthesia dose complications. In addition, the
reduced surgical time could reduce the overall surgical costs due
to a reduction in the time spent in the operating room.
[0123] Another benefit of eliminating the surgical fit is increased
consistency. Surgically adjusting and fitting each device tends to
introduce variability between patients by any given surgeon. In
addition, there is variability between different surgeons and
hospitals that can only be reduced by rigorous implant training.
Thus, eliminating the surgical fit procedure may reduce the
variability and the training requirements.
[0124] Eliminating the surgical fit can also make implanting the
device more compatible with minimally invasive surgical approaches.
Typically, the surgical fit step for implant of the original device
requires access to the anterior portion of the heart. This is most
commonly accomplished using a full median sternotomy surgical
approach. If a high compliance device can allow the appropriate fit
to be obtained through device size selection rather than surgical
customization of fit, the surgery may be possible through a smaller
incision than a full median sternotomy. It may even be possible to
implant the device through small portal incisions.
[0125] Minimally invasive surgical incisions can have numerous
benefits, including reduced pain, less cosmetic scaring, faster
hospital release and faster return to physical activities. The
reduced hospital stay from minimally invasive surgery can also help
to reduce overall surgical costs and make the surgery more accepted
and routine in the medical community.
[0126] In one embodiment, the cardiac support device is
3-dimensional shape that is constructed from a flat fabric mesh. To
form a 3-dimensional shape from a flat fabric, the device typically
includes sewn seams where the device material is a little denser
and thicker. Unfortunately, the seams can result in a greater
tissue response and an increased potential for adhesion between the
device and adjacent tissues in the chest, other than the heart.
[0127] Thus, in another embodiment, the device is manufactured
using advanced fabrication methods that eliminate the manufactured
seams. However, even if the pre-fabricated seams can be eliminated,
the surgical fitting may result in a seam that is even thicker and
denser than those produced during device manufacture. Consequently,
if the surgical seam from fitting is eliminated, for example, by
using a high compliance device, and the manufactured seams are
eliminated, the potential for adhesions to adjacent tissues would
be reduced. This could be important, particularly when future
surgeries require access to the chest cavity. Adhesions make
surgical access much more difficult.
[0128] 3. More Volume Overloading Tolerant
[0129] Another advantage of a high compliance device is the ability
of the device to expand and not over restrain the heart in the case
of volume overloading. For example, excessive fluid intake can
impact the volume of the heart. A high compliance device may
benefit the patient by helping to support the increased volume
loading, without overly restricting the heart as might occur with
the cardiac condition known as constriction. Although dilatation
constraint is a potentially important mechanism of the cardiac
support device, higher compliance may provide:adequate support and
resistance to dilatation without overly restricting the patient's
normal variations in fluid intake.
[0130] Having disclosed the invention in a preferred embodiment
modifications and equivalents will become apparent to those skilled
in the art. It is intended such modifications and equivalents shall
be included within the scope of the appended claims. For example,
while the invention is described covering the ventricles, the
invention can cover one or both of the atria only or in combination
with ventricle coverage. Also, the device can be provided with
circumferential fibers which have a maximum stretch (or no stretch)
at a volume representing a maximum volume for end diastole at time
of placement. Such a modification provides acute prevention of
diastolic expansion beyond a maximum. Use of multiple sets of
fibers are described in Haindl international patent application PCT
WO 98/58598 published Dec. 30, 1998.
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