U.S. patent application number 10/318884 was filed with the patent office on 2003-05-08 for passive cardiac assistance device.
This patent application is currently assigned to ABIOMED, INC.. Invention is credited to Kung, Robert T.V., Lederman, David M., Rosenberg, Meir.
Application Number | 20030088151 10/318884 |
Document ID | / |
Family ID | 27049933 |
Filed Date | 2003-05-08 |
United States Patent
Application |
20030088151 |
Kind Code |
A1 |
Kung, Robert T.V. ; et
al. |
May 8, 2003 |
Passive cardiac assistance device
Abstract
Artificial implantable active and passive girdles include a
heart assist system with an artificial myocardium employing a
number of flexible, non-distensible tubes with the walls along
their long axes connected in series to form a cuff and a passive
girdle is wrapped around a heart muscle which has dilatation of a
ventricle to conform to the size and shape of the heart and to
constrain the dilatation during diastole. The passive girdle is
formed of a material and structure that does not expand away from
the heart but may, over an extended period of time be decreased in
size as dilatation decreases.
Inventors: |
Kung, Robert T.V.; (Andover,
MA) ; Lederman, David M.; (Marblehead, MA) ;
Rosenberg, Meir; (Newton, MA) |
Correspondence
Address: |
NUTTER MCCLENNEN & FISH LLP
WORLD TRADE CENTER WEST
155 SEAPORT BOULEVARD
BOSTON
MA
02210-2604
US
|
Assignee: |
ABIOMED, INC.
Danvers
MA
|
Family ID: |
27049933 |
Appl. No.: |
10/318884 |
Filed: |
December 13, 2002 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
10318884 |
Dec 13, 2002 |
|
|
|
09223645 |
Dec 30, 1998 |
|
|
|
6508756 |
|
|
|
|
09223645 |
Dec 30, 1998 |
|
|
|
09023592 |
Feb 13, 1998 |
|
|
|
6224540 |
|
|
|
|
09023592 |
Feb 13, 1998 |
|
|
|
08581051 |
Dec 29, 1995 |
|
|
|
5800528 |
|
|
|
|
08581051 |
Dec 29, 1995 |
|
|
|
08490080 |
Jun 13, 1995 |
|
|
|
5713954 |
|
|
|
|
Current U.S.
Class: |
600/37 |
Current CPC
Class: |
A61F 2002/0068 20130101;
A61F 2/2481 20130101; A61M 60/40 20210101; A61M 2205/33 20130101;
A61M 60/122 20210101; Y10S 623/904 20130101; A61M 60/869 20210101;
A61M 2205/3303 20130101; A61M 60/857 20210101; A61M 60/50 20210101;
A61M 60/268 20210101 |
Class at
Publication: |
600/37 |
International
Class: |
A61F 002/02 |
Claims
1. A method for treatment of a patient, whose heart is
characterized by ventricular dilatation comprising the steps of,
wrapping a girdle around at least the ventricle of said patient's
heart; and maintaining said girdle in a passive state for an
extended period of time, said girdle being formed such that it can
adjust in size and shape to conform to the outer shape of said
ventricle and to not expand its dimension in a direction away from
said natural heart.
2. A method in accordance with claim 1 wherein said girdle is
formed of a sheet of material prestressed in the plane of said
sheet to a value below the elastic limit of said material, said
sheet having a tension which limits extension away from said heart,
while providing compression forces radially inward toward said
heart.
3. A method in accordance with claim 1 wherein said girdle is
formed of a vertically oriented series of sealed, independent,
generally horizontally extended cylindrical segments, said vertical
orientation being parallel with an axis of the heart running to its
apex, including the further step of, introducing fluid into said
cylindrical segments to decrease the inner perimeter of said girdle
so that its size conforms generally to the size of said patient's
ventricle.
4. A method in accordance with claim 3 and further including means
for increasing the volume of fluid within said sealed volume in a
controlled fashion to decrease the dimensions of the inner lining
when said ventricle decreases in size over an extended period.
5. A method in accordance with claim 1 wherein said girdle is
formed of a net of interlocking loops, unattached to one
another.
6. A method in accordance with claim 5, wherein and said
interlocked loops are further interconnected by strings extending
in at least a first dimension and including the further step of
pulling on said strings to decrease the dimension of said girdle in
position around said patient's heart.
7. A method in accordance with claim 3 and further including the
step of, placing an inner lining between said cylindrical segments
and the outer surface of said patient's heart and, placing a
tension measuring sensor at the interface between the outer surface
of said patient's heart and the inner lining of said girdle, said
sensor providing an output signal indicative of the tension of said
lining adjacent to said sensor, and, providing said output signal
to means for adjusting the amount of fluid within said cylindrical
segments until the tension at said lining is at a predetermined
value.
8. Apparatus for providing passive support to a natural heart
characterized by ventricular dilatation comprising, a girdle for
wrapping around said heart, said girdle being formed of a sheet of
material prestressed in the plane of said sheet to a value below
the elastic limit of said material, such that said sheet has a
tension which limits extension away from said heart, while
providing compression forces radially inward toward said heart.
9. Apparatus for providing passive support to a natural heart
characterized by ventricular dilatation comprising, a girdle for
wrapping around said heart, said girdle comprising a net of
interlocked plastic loops, not attached to one another.
10. Apparatus for providing support to a heart characterized by
dilatation of a ventricle comprising, a girdle formed of a
vertically oriented series of sealed, independent, generally
horizontally oriented cylindrical segments, said vertical
orientation being parallel with an axis of the heart running to its
apex, means for providing a fluid to said segments to control the
dimensions of said girdle, and means for changing the volume of
fluid provided to said sealed volume only after an extended period
of time.
11. Apparatus in accordance with claim 10 and further including a
distensible lining formed between said cylindrical segments and
said natural heart.
12. Apparatus in accordance with claim 10 and further including
means for providing as an interface between said inner lining and
the outer surface of said heart, a cellular wall constructed of
myocardial cells harvested from said heart and mounted on a
scaffold to generate said cellular wall.
13. Apparatus for providing support to a heart characterized by
dilatation of a ventricle, comprising a girdle formed with open
ends, the wall of said girdle being formed of an interconnected
vertically oriented series of closed, horizontally oriented, tubes
constructed of a flexible, nondistensable material, forming a
fluidically sealed volume, said vertical orientation being aligned
with an axis of the heart running to its apex, means for
controllably filling said tubes with a fluid, such that when
completely filled, each of said tubes assumes a tubular cylindrical
shape with a circular cross section, and when emptied of said
fluid, each of said tubes assumes a collapsed shape, and means for
providing a selected amount of filling to said tubes with said
fluid, and changing said amount only in response to decrease of
said dilatation after an extended period of time.
14. A method of generating the interface between the interior of an
external girdle for a natural heart and the myocardium of said
natural heart comprising the step of, providing a scaffold of
biologically inert material between the lining of said girdle and
the exterior myocardial surface of said heart and generating a wall
on said scaffold by application of myocardial cells to said
scaffold, said myocardial cells being harvested from said heart.
Description
[0001] This application is a continuation of U.S. patent
application Ser. No. 09/023,592 filed Feb. 13, 1998, which is a
divisional of U.S. patent application Ser. No. 08/581,051 filed
Dec. 29, 1995 (now U.S. Pat. No. 5,800,528 issued Sep. 1, 1998),
which is a continuation-in-part of U.S. patent application Ser. No.
08/490,080 filed Jun. 13, 1995 (now U.S. Pat. No. 5,713,954 issued
Feb. 3, 1998). The contents of each of these patent applications
are specifically incorporated herein by reference.
FIELD AND BACKGROUND OF THE INVENTION
[0002] The present invention relates to cardiac devices in general,
and more specifically to passive and active cardiac girdles.
[0003] Patients having a heart condition known as ventricular
dilatation are in a clinically dangerous condition when the
patients are in an end stage cardiac failure pattern. The
ventricular dilatation increases the load on the heart (that is, it
increases the oxygen consumption by the heart), while at the same
time decreasing cardiac efficiency. A significant fraction of
patients in congestive heart failure, including those who are not
in immediate danger of death, lead very limited lives. This
dilatation condition does not respond to current pharmacological
treatment. A small amount, typically less than 10%, of the energy
and oxygen consumed by the heart, is used to do mechanical work.
Thus the balance, which is the major part of the energy consumed by
the heart is used in maintaining the elastic tension of the heart
muscles for a period of time. With a given pressure, the elastic
tension is directly proportional to the radius of curvature of the
heart ventricle. During ventricular dilatation the ventricular
radius increases and the energy dissipated by the heart muscle just
to maintain this elastic tension during diastole is abnormally
increased, thereby increasing oxygen consumption.
[0004] A number of methods and devices have been employed to aid
the pumping action of failing hearts. Many of these include sacs or
wraps placed around the ailing heart, or, in some instances only
around the ventricle of the failing heart, with these wraps
constructed to provide for active pumping usually, but not always,
in synchronism with the ventricular pumping of the natural heart. A
number of cardiac assist systems employing a variety of pumping
approaches for assisting the pumping action of a failing natural
heart have been developed. These systems include those suitable for
partial to full support of the natural heart, short term (a few
days) to long term (years), continuous pumping to various degrees
of pulsability, and blood contacting versus non-blood contacting.
Table 1 lists a number of presently developed devices with
pertinent operating characteristics.
1TABLE 1 Level of Blood Device Support Pulsatility Duration
Contacting Comments IAPB Partial <20% Y Days to Months Y
Counterpulsation provides LV unloading Biopump Full N Days Y
Limited to short duration due to thrombotic potential Thoratec Full
Y Months Y Sac-type actuation Novacor Full Y Months Y Sac-type pump
with electric actuation Hemopump Partial N Days Y Axial flow pump
50-75% Heart Mate Full Y Months Y Pusher-Plate pneumatic and
electric Aortic Patch Partial Y Months Y Counterpulsation BVS 5000
Full Y Weeks Y Designed for temporary support Anstadt Full Y Days N
Cardiac resuscitation Cardiomyoplasty Partial <20% Y Years N
Requires muscle training for active support
[0005] One, more recent development in the field of cardiomyoplasty
involves the wrapping and pacing of a skeletal muscle around the
heart to aid in the pumping. In that configuration, a pacemaker is
implanted -to control the timing of the activation of the wrapped
around skeletal muscle.
[0006] A major consideration in the design of cardiac support
systems is the risk of thromboembolism. This risk is most
associated with use of artificial blood contacting surfaces. A
variety of approaches have been employed to reduce or eliminate
this problem. One approach has been the employment of smooth
surfaces to eliminate potential sites for thrombi and emboli
generation as well as textured surfaces to promote cell growth and
stabilization of biologic surfaces. One problem affecting
thromboembolism risk in heart assists arises from the use of
prosthetic, biologic or mechanical pericardial valves. This risk
can be some what lowered by the use of anticoagulation therapy.
However, the use requires careful manipulation of the coagulation
system to maintain an acceptable balance between bleeding and
thromboembolic complications. The textured surface approach employs
textured polyurethane surfaces and porcine valves to promote
pseudo-intima formation with a stable cellular lining. While
thromboembolic rates resulting from these measures are acceptable
as temporary measures, improvements, particularly for implantable
devices are highly desirable.
[0007] A second problem associated with implanted cardiac assist
devices is the problem of infection, particularly where the
implanted device has large areas of material in contact with blood
and tissue. More recently clinical protocols have improved and even
the drive line and vent tubes associated with implants that require
some percutaneous attachments have been manageable. However, for a
ventricular assist device, quality of life considerations require
that vent lines and drive lines which cross the skin barrier be
eliminated thereby avoiding the encumbrance to patient
activities.
[0008] A third problem area in ventricular assist devices is the
calcification of these devices. This is particularly so for long
term implant situations which may last five years or more. Here
again the criticality of this factor is reduced for devices which
do not involve direct blood contact.
[0009] Another approach employed in ventricular assist has been the
development of non-pulsatile pumps. However, once again, the blood
is exposed to the surfaces of the pump, particularly the bearing
and seal area.
[0010] Unlike an entirely artificial heart, in which failure of the
system leads to death, a ventricular assist device augments the
impaired heart and stoppages should not result in death, unless the
heart is in complete failure. However, for most present ventricular
assist device systems, stoppage of even a few minutes results in
formation of blood clots in the device, rendering any restart of
the system a very risky undertaking.
SUMMARY AND OBJECTS OF THE INVENTION
[0011] According to one aspect, the present invention an artificial
myocardium is constructed of an extremely pliant, non-distensible
and thin material which can be wrapped around the ventricles of a
natural, but diseased heart. This artificial myocardium mimics the
contraction-relaxation characteristics of the natural myocardium
and provides sufficient contractility, when actuated, to at least
equal the contractility of a healthy natural myocardium. In this
arrangement all of the direct blood contact is with the interior
surfaces of the natural heart and surrounding blood vessel system.
The device is hydraulically actuated in timed relationship to the
contractions of the natural heart.
[0012] Using this system, the natural heart is left in place and
the assist system supplies the reinforcing contractile forces
required for satisfactory ventricular ejection.
[0013] A key concept for this artificial myocardium system is
achieved by the realization of a controllable, artificial
myocardium employing a cuff formed of a series of closed tubes
connected along their axially extending walls. With sufficient
hardware to hydraulically (or pneumatically) inflate and deflate
these tubes, a controlled contraction is produced as a result of
the geometric relationship between the length of these series of
tubes in deflated condition and the length of the series of tubes
when they are fluidically filled in the inflated condition. If the
cuff is formed of a series of "n" tubes, each of diameter "d" when
inflated, connected in series, the total perimeter length of this
cuff when deflated is given by n(.pi.d/2). However, when these
tubes are filled with fluid, they have a circular cross-section
such that the length of the cuff is the sum of the diameters in the
individual tubes or nd. Thus the ratio of the change in perimeter
length between the collapsed and the filled state is .pi./2. If
this cuff is wrapped around the natural heart, it will, when
pressurized, shorten and squeeze the heart by producing a
"diastolic" to "systolic" length change of 36%. Typical sarcomere
length changes are approximately 20%.
[0014] Suitable hardware, including a hydraulic pump, a compliant
reservoir and rotary mechanical valve, together with appropriate
actuating electronics can all be implanted in the patient's body.
If the power source is an internal battery, then power may be
transcutaneously transmitted into the body to recharge this
battery.
[0015] Ventricular dilatation is a clinically dangerous condition
for end stage cardiac failure patients. The output of the heart is
effected by: (a) end-diastolic volume (ventricular volume at the
end of the filling phase), (b) end-systolic volume (ventricular
volume at the end of the ejection phase), and (c) heart rate. When
(a) is very large, (b) also tends to be larger and (c) tends to be
larger than normal. All three of these factors contribute to large
increases in the tension-time integral and therefore to increased
oxygen consumption.
[0016] Only a small amount of the energy consumed by the heart is
used to do mechanical work. For example, with a cardiac output of 5
liters/minute, and .DELTA.p of 100 mm(Hg), the mechanical work done
by the left ventricle is about 1.1 watts, and that of the right
ventricle is about 0.2 watts. This compares with the typical total
energy consumed by the heart (mechanical work during systole plus
the energy cost in maintaining elastic tension during diastole) of
about 12 to 15 watts.
[0017] Thus, since cardiac efficiency (typically between 3% and
15%) is defined as the ratio of the mechanical work done by the
heart to the total energy (or load of the heart muscle): then,
[0018] Cardiac Efficiency, 1 = P v V P V + k T r
[0019] where
[0020] P.sub.v: Ventricle Pressure
[0021] P: Pressure
[0022] V: Volume
[0023] T: Tension
[0024] t: Time
[0025] The constant k accounts for conversion of units.
[0026] An increase in mechanical work by a large factor results in
a small increase in oxygen consumption but an increase in tension
time causes a large increase in oxygen consumption. Patients with
dilated ventricles who have undergone active cardiomyoplasty have
not been reported to show any objectively measurable hemodynamic
improvement.
[0027] According to a further aspect of the present invention, a
completely passive girdle is wrapped around the ventricle or the
entire heart muscle, and sized so that it constrains the dilatation
during diastole and does not effect the action of the ventricle
during systole. With the present surgical techniques, it is
expected initial access to the heart to place the girdle in
position, will require opening the chest. However, it may be
possible to locate a girdle in position without thoracotomy. In one
embodiment, a synthetic girdle made from material that can limit
tension, but is otherwise deformable to conform to the anatomical
geometry of the recipient heart is used. This girdle may be
adjustable in size and shape over an extended period of time in
order to gradually decrease the ventricular dilatation. A second
embodiment employs a fluid filled passive wrap constructed of a
series of horizontal sections. This provides for a variable volume
to be enclosed by the wrap with volume control being obtained by
controlling the volume of fluid from an implantable reservoir
within the body. In its most preferable form, this passive wrap can
be formed of a series of horizontal tubular segments each
individually sealed and attached to one another along the long axis
of the cylinder. If the cylinders are made of indistensible
material, then changing the volume of fluid from the cylinders
being in a substantially deflated condition to one where they are
partially or fully inflated, decreases the internal perimeter of
the wrap or girdle, thereby decreasing the effective radius of the
girdle around the heart. Another feature of the invention is a
feedback system, wherein sensors, for example, strain gauges, can
be built into an indistensible lining to measure its tension and
thereby provide automatic feedback to a hydraulic circuit
controlling the wrap volume.
[0028] To avoid the problem of potential irritability and damage to
the external myocardium cells by virtue of the artificial wrap and
its long term constraining contact with the myocardium, one
embodiment of the invention employs a tissue engineered lining to
protect the myocardium. This tissue engineered lining consists of a
polymer scaffold seeded with myocardial cells harvested from the
patient's own myocardium using tissue engineering technology. That
lining then generates a biological myocardio-interfacing surface
and remains firmly attached to the polymer interfacing with the
surface from which the wrap is made. Such a lining would integrate
biologically to the heart's myocardial cells in a manner analogous
to other devices currently being investigated which use cell
scaffolds for in vitro and in vitro tissue engineering.
[0029] It is therefore an object of the present invention to
produce a ventricular assist device system employing an artificial
myocardium placed around the natural heart (extra cardiac assist).
This design, then, does not contact the bloodstream eliminating
many of the problems discussed above.
[0030] It is another object of this invention to provide a
ventricular assist device which mimics the action of the natural
heart while avoiding the compressive action of the direct
mechanical ventricular actuation systems on the epicardium.
[0031] It is a further object of this invention to provide an
artificial myocardium in which the external fluid being pumped is a
fraction of the blood volume pumped by the action of the artificial
myocardium.
[0032] It is yet another object of this invention to provide a
ventricular assist device which is compact, requires relatively low
energy input and does not require percutaneous components.
[0033] It is a further object of this invention to provide a
passive girdle to be wrapped around a heart suffering from
ventricular dilatation to limit this dilatation and thus improve
the performance characteristics of the heart.
[0034] It is another object of this invention to provide a passive
girdle or vest which can, over a period of time, have its diameter
decreased to effect some decrease in dilatation of the
ventricle.
[0035] Other objects will become apparent in accordance with the
description of the preferred embodiments below.
BRIEF DESCRIPTION OF THE DRAWING
[0036] FIGS. 1A and 1B are diagrammatic illustrations of the tube
construction of an artificial myocardium in accordance with the
principles of this invention;
[0037] FIG. 2 is a generally diagrammatic illustration of the
anatomical placement of the cardiac assist system of this invention
in the human body;
[0038] FIG. 3A is an illustration in diagrammatic form of the
perimeter length of the artificial cuff at various stages of the
natural heart contractions;
[0039] FIG. 3B is an illustration of a cross-sectional view of the
systolic and diastolic shapes of the artificial myocardium;
[0040] FIG. 4A is a diagrammatic view of a partial wraparound of
the left ventricle by the artificial myocardium in the systolic
state and FIG. 4B is a diagrammatic view of a partial wraparound of
the left ventricle by the artificial myocardium in the diastolic
state;
[0041] FIG. 5 is a diagrammatic illustration of the geometric
relationship of the partially inflated tubes and the encircled
heart represented by a radius of D/2.
[0042] FIGS. 6A, 6B and 6C illustrate factors involved in altering
the length of the individual tubes when inflated or deflated, to
achieve specific shrink ratios.
[0043] FIG. 7 is a graphical representation of the ratio of
hydraulic to afterload pressure as a function of the tube inflation
parameter, .theta., for n=10;
[0044] FIG. 8 is a graphical representation of the timing
relationship for device actuation;
[0045] FIG. 9 is a graphical representation of the left ventricular
pressure versus the left ventricular volume under various
conditions;
[0046] FIG. 10 is a block diagram of the hydraulic system of the
artificial myocardium of this invention;
[0047] FIG. 11A is a cross-sectional view of an energy converter
and valve structure in one position for use in the artificial
myocardium of this invention;
[0048] FIG. 11B is a cross sectional view of the same structure as
FIG. 11A but with the valving in a different position for pumping
from the hydraulic cuff to the hydraulic reservoir;
[0049] FIG. 12A is a cross-sectional view across the skin interface
of a subdermal port for emergency access and manual pumping;
[0050] FIG. 12B is an "X-ray" view of the access system of FIG. 12A
viewed externally;
[0051] FIG. 13 is a graphical representation of the hydraulic to
physiological pressure ratios as a function of the number of tubes
for both univentricular and biventricular support;
[0052] FIG. 14 is a graphical representation of the physiological
to hydraulic volume ratios as a function of the number of tubes for
both univentricular and biventricular support;
[0053] FIG. 15 is a graphical representation of the blood stroke
volume as a function of the number of tube segments in the
artificial myocardium of this invention;
[0054] FIG. 16 is a diagrammatic illustration of a mock loop used
for in vitro tests;
[0055] FIG. 17 is a graphical representation of the relationship
between Afterload Pressure (AOP) and Driving Pressures (P-Drive)
for the artificial myocardium; and
[0056] FIG. 18 is a graphical representation of flow sensitivity to
drive pressure at a constant afterload pressure;
[0057] FIG. 19A is an illustration generally in cross sectional
form of a heart girdle constructed in accordance with the
principles of this invention;
[0058] FIG. 19B is an illustration in cross-sectional form of the
heart girdle of FIG. 19A with the girdle in a pneumatically filled
condition;
[0059] FIG. 20 is a perspective view of the heart girdle of FIGS.
19A and 19B showing the horizontal segments.
[0060] FIG. 21 is an illustration generally in block diagram form
of a control system for the heart girdle of FIG. 19 including a
strain gauge and electronic actuator to maintain constant tension
at the interface between the girdle and the heart muscle;
[0061] FIG. 22 is an illustration in perspective view of a heart
girdle employing a flexible mesh of interlocked circular plastic
loops;
[0062] FIG. 23 is an illustration of a portion of a girdle
constructed generally in accordance with the girdle construction of
FIG. 22, but further including strings adapted to draw the girdle
into decreasing diameter shape;
[0063] FIG. 24 is another embodiment of a portion of a passive
girdle formed of a material characterized by a specific internal
structure; and
[0064] FIG. 25 is a cross sectional drawing of a girdle-myocardium
interface constructed of biologically engineered myocardial
tissue.
DESCRIPTION OF PREFERRED EMBODIMENTS
[0065] FIG. 1A and FIG. 1B illustrate diagrammatically the
operation of the artificial myocardium. The artificial myocardium
11 is formed of a series of tubes planed together in series to
form, in this instance a complete circle, which in FIG. 1A has a
diameter D.sub.d. In FIG. 1B the feature of the tubes is filled
hydraulic fluid producing a circular cross-section, shortening the
total perimeter of the circular cuff to a circle having a diameter
D.sub.S. Referring to FIG. 1B, if the diameter of the tube with the
circular cross-section is d, then the diameter of the circular cuff
is approximately equal to nd/.pi., where n equals the total number
of tubes. On the other hand, when the tubes are no longer filled
with hydraulic fluid and are collapsed then the diameter D.sub.d is
approximately equal to 2 n ( d / 2 )
[0066] These expressions follow from the consideration that the
series of n tubes in the inflated condition, as illustrated in FIG.
1B form a circle with the number of tubes times diameter of each of
the individual tubes. On the other hand in the collapsed condition
each one of the tubes has a length equal to its perimeter divided
by 2. Since the perimeter is .pi.d then the length of each
collapsed tube is .pi.d /2 and the diameter of the cuff in this
condition is the sum of the length of the collapsed tubes over
.pi..
[0067] When this cuff is placed around a natural heart and the
filling and emptying of the tubes is in phase with the systole and
diastole of the natural heart, then the shortening of the cuff
forces the surrounded ventricle to decrease its diameter thereby
causing the ventricle to eject blood. The ejection fraction of this
artificial myocardium is independent of the number of tubes or the
heart dimensions. The ejection fraction is a function of only the
hydraulic pressure. When the hydraulic pressure is large enough to
inflate the tubes to cylinders, the ejection fraction is, 3 E f = D
d 2 - D S 2 ( D d - 2 t ) 2 where D S nd , D d = nd 2 , t = heart
muscle thickness
[0068] FIG. 2 illustrates an artificial myocardial assist system
located in the human body. In the illustrated system the artificial
myocardium 11 is shown as a cuff placed around the ventricles of
the natural heart 17. The hydraulic fluid is pressurized by energy
converter 19 either in the direction of the cuff 11 or of a
compliant hydraulic reservoir 21. The energy converter 19 is
electrically controlled by virtue of internal electric circuit 23
which is powered by an internal battery 25. The internal electrical
circuit 23 is also coupled to external battery 27 via a
transcutaneous electrical terminal (TET) 31. The energy converter
19 consists of a hydraulic pump coupled to a brushless electric
motor to shuttle fluid between the artificial myocardium 11 and the
compliant reservoir 21. Flow switching is accomplished by a rotary
mechanical valve incorporated into the energy converter, which in
turn is synchronized by a control signal generated by detection of
the R wave from the ECG signal in the natural heart. Continuous
adjustment of the hydraulic pump output allows the level of cardiac
assist to be varied on a beat-by-beat basis.
[0069] FIG. 3A is a graphical illustration of the length of the
perimeter of the artificial myocardium during systole, mid systole
and diastole. FIG. 3B is an illustration in cross section view of
the systolic and diastolic shapes of the artificial myocardium in a
cylindrical geometry. The outer ring illustrates the cylindrical
tubes in collapsed form, while the inner ring illustrates those
same tubes when they are filled with hydraulic fluid during the
systole. The natural heart pumps blood primarily through
circumferential contraction. Most of the diastolic to systolic
volume change is derived primarily from the 20% change in the
circumference component and to a lesser extent the 9% change in the
axial length. As can be seen in FIG. 3A and FIG. 3B, the volumetric
change of the myocardium is 36% from the relaxed (diastole)
position to the fully contracted (systole) position. In the
cross-sectional view, and assuming that the artificial myocardium
were a completely cylindrical cuff, there is a 60% change in the
area between these states in the artificial myocardium, equivalent
to a 60% ejection fraction of a healthy heart. Although the
description is based on a cylindrical geometry, with
interconnecting tapered tubes, the artificial myocardium will match
the conical shape of the heart when appropriate taper angle is
selected for the tubes.
[0070] With this hydraulic design, the natural heart having a
typical myocardium thickness, a heart base diameter of 80 mm and an
axial ventricular length from apex to base of 50 mm, a left
ventricular wrap of the artificial myocardium results in a stroke
volume of 83 cc. These values are the same as that which would be
expected from a normally operating left ventricle.
[0071] One very important factor in the operation of the artificial
myocardium is that the hydraulic pressure required for contraction
against a given ventricular pressure is directly proportional to
the number of tubes n forming this artificial myocardium. From
energy conservation principles, the hydraulic flow in this
artificial myocardium varies inversely with the number of tubes.
For example, with typical natural heart dimensions, and a hydraulic
stroke volume of 24 cc, a pressure of 760 mm Hg produces a left
ventricular stroke volume of 95 cc at a mean aortic pressure of 90
mm Hg. Importantly, the hydraulic flows required are much less than
the generated blood flow. In this example, the hydraulic flow is
approximately 25% of the blood flow produced. These smaller
hydraulic flows result in lower hydraulic losses and higher
efficiencies. This, taken together with the smaller dimensions for
the energy converter and the compliance chamber is very
advantageous for an implanted device.
[0072] The output of a single energy converter can simultaneously
generate different contractile forces for left and right ventricle
assist by varying the number of tubes which are wrapped around the
left and right ventricles while maintaining an equal drive
pressure. If the artificial myocardium on the right side has M
times the number of tubes as that on the left side, the contractile
pressures on the right side will be M times lower. With this
arrangement, the artificial myocardium may be tailored to match
differing afterloads from the two ventricles. A good design
parameter for considerations for efficiency and tube dynamics would
be to operate at a contraction of 22%, which is very close to the
contraction value of a natural healthy myocardium. It is believed
that with this contraction level, the low mechanical stresses on
the artificial myocardium may well result in an operation life of
five years, a high reliability for the artificial myocardium.
[0073] In the artificial myocardium assist system of this
embodiment the contribution from the artificial myocardium is
additive to the natural heart with a timing cycle synchronized with
the ECG, so that the control algorithm can adjust the hydraulic
flow on a beat-by-beat basis to achieve the desired ejection
fraction. Thus, if the natural myocardium was completely healthy,
minimal pumping would be required by the hydraulic pump. On the
other hand, if the natural heart had very little myocardial
contractility, then the artificial myocardium would provide almost
the entire contracting force.
[0074] With this design of the artificial ventricles there is no
blood contact with the artificial surfaces of the cardiac assist
system, thereby avoiding the principal concerns of the
thromboembolism risk. Another important safety consideration, is
that if the artificial myocardium system were to stop, the natural
forces on the hydraulic fluid will cause it to empty from the
artificial myocardium and flow into the compliant reservoir. The
only effect of these conditions on the cardiovascular system would
be those caused by the collapsed flexible wrap on the myocardium. A
subdermal port could be provided to allow emergency actuation of
the artificial myocardium with a pneumatic pump placed external to
the patient's body. Another consideration is that the control
algorithm of the artificial myocardium assist system can be
arranged to provide contractions at a fixed predetermined rate if
there should be a ventricular fibrillation or tachycardia of the
natural heart.
[0075] Because the design provides for no contact between the blood
and the components of the artificial myocardium assist system, the
biocomparability factors are limited to those which relate to the
interface between tissues and these components. In this embodiment,
the tissue contacting material of the artificial myocardium may be
a polyetherurethane, Angioflex.RTM., manufactured by ABIOMED, Inc.
of Danvers, Mass.
[0076] FIGS. 4A and 4B illustrate an example of a partial wrap
around the heart for left ventricular support. The full wrap can be
used for biventricular support. FIGS. 4A and 4B show the systolic
and diastolic positions of the artificial myocardium under this
condition. Each of the n tubes are attached at their outer wall to
two neighboring tubes except for the two ends. Each tube, when
inflated, has a diameter d, resulting in a wrap length of nd.
Conversely, when the tubes are deflated the wrap length is
n.pi.d/2
[0077] As illustrated in FIG. 5, when the tubes are partially
inflated the length of the wrap is given by
L=nd.pi.sin.theta./,2.theta.
[0078] where .theta. is the angle representative of the curvature
of the wall of the partially inflated tube and is defined as
.theta..pi.d/4r
[0079] where r is the radius of the arc of each half of the tube
wall when inflated.
[0080] Although the ventricle is conical in shape and accordingly
the artificial myocardium is conformed to that shape, for
simplicity the representation in FIGS. 4A and 4B is of a
cylindrical shape. In FIG. 3B the effects of both tube inflation
and stroke volume are shown. The systolic contracted shape 13a of
the artificial ventricle is plotted concentrically inside the
diastolic distended shape 13b. The shaded annulus portion 15 of
FIG. 3B represents the stroke volume change due to contraction,
while the dotted circle enclosed portion of the inflated tubes
represents displacement volume change. For the artificial
myocardium this displacement volume has a minor contribution to the
stroke volume.
[0081] The volume may be expressed as 4 S V = 4 D d 2 [ 1 - ( 1 - +
sin ) 2 ] l + 1 2 V H
[0082] where l is the perimeter length of the cuff, where D.sub.d
is, as shown, the diastolic diameter of the ventricle and .chi. is
the fraction of the ventricle being wrapped. V.sub.H represents the
actual displacement due to the effects of tube inflation in
addition to the stroke volume derived from the contraction. If
values are substituted in this equation, assuming a typical natural
heart diameter of approximately 8 cm, a length approximately 5 cm
and a left side partial wrap of .chi.=1/2, the contractile change
from the uninflated to fully inflated tubes results in a stroke
volume of approximately 83 cc. V.sub.H is assumed to be zero.
Similarly for a full wrap, the stroke volume would be approximately
150 cc, equivalent to the sum of left and right side stroke
volumes.
[0083] The ejection fraction achievable using such an assist device
may be estimated. The diastolic ventricular volume is given by 5 V
d = 4 ( D d - 2 t d ) 2 l ,
[0084] where t.sub.d is the ventricular wall thickness. Since the
ejection fraction is defined as the stroke volume divided by the
diastolic ventricular volume, it is given by 6 EF = SV 4 ( D d - 2
t d ) 2 l ,
[0085] Thus, the maximum ejection fraction obtainable for a typical
wall thickness of 1 cm and a D.sub.d equal to 8 cm is 59%, a number
which is consistent with the ejection fraction for a normal healthy
heart.
[0086] FIG. 6A describes a section of a tube having an inflated
diameter d and an inflated length l. The tube ends are formed to
hemispheres in the inflated position, the hemisphere radius being
d/2.
[0087] The following relation can be written:
l.sub.i=l.sub.o+d (1)
[0088] where:
[0089] l.sub.i is the inflated length of the tube
[0090] l.sub.o is the straight part of the tube
[0091] d is the inflated diameter of the tube.
[0092] FIG. 6B describes the tube when it is deflated. The
flattened diameter of the tube is Bd/2 and its flattened length
is:
l.sub.d=l.sub.o+Bd/2 (2)
[0093] where:
[0094] l.sub.d is the deflated length of the tube.
[0095] The shrink ratio, defined as the change in linear dimensions
relative to the original linear dimension l.sub.d, is: 7 R = [ l o
+ d 2 ] - [ l o + d ] l o + d 2 = d [ 2 - 1 ] l o + d 2 = - 2 2 l o
d + ( 3 )
[0096] Where:
[0097] R=relative change in linear dimensions.
[0098] As can be seen from the equation, the ratio R is only a
function of the ratio between the diameter d and the length
l.sub.o. It can also be shown that this ratio does not change if a
number of these tubes are connected in series provided l.sub.o/d
remains unchanged.
[0099] Example: If the required longitudinal shrinkage is 12%, the
ratio l.sub.o/d can be calculated, using equation (3) to be: 8 0.12
= - 2 2 l o + d
[0100] To achieve a contraction of 12% longitudinally, the tube
length l.sub.o should be 3.18 times the diameter d. In a particular
case, where the tube diameter is 10 mm, the overall tube length
will be 41.8 mm. This length is about half of the required length
and therefore, two of these tubes will be connected in series to
achieve this goal. The series connection can be done by making two
wraps of half length, or, making the individual tubes with a shape
as described in FIG. 6C.
[0101] As discussed earlier the contractile action of the
artificial myocardium results from the inflation of the series of
tubes that are physically attached to each other. The inflation of
these tubes can be accomplished either pneumatically or
hydraulically. For a permanently implantable device, the hydraulic
approach is more practical. With a pneumatic system, even in the
absence of leakage losses, gas permeation across flexing membranes
is unavoidable. This effect is not probable in a hydraulic system
with a proper choice of working fluids. In addition, the hydraulic
system is safer in the event of rupture failure since high pressure
cannot be maintained in the event of a leak. In the case of a
pneumatic system, a severe leak could result in cardiac
compression. In the situation where the artificial myocardium is
being employed as an assist device, the natural myocardium can
generate some tension, and the artificial myocardium need only
generate sufficient tension to boost the intraventricular pressure.
Accordingly, the differential pressure, P required of the
artificial myocardium may be only 20-30 mm Hg, boosting the
ventricular pressure from, for example, 60 to 80-90 mm Hg. The
ratio of hydraulic pressure P.sub.H to the load pressure P is
expressed by 9 P H P = n tan 2 cos + 1 2 ( 1 - tan tan )
[0102] where .alpha. is the half angle subtended by each tube
centered to the ventricular axis.
[0103] Accordingly, the pump pressure P.sub.H is related to three
parameters. First it is directly proportional to load pressure.
Thus higher load pressures requires higher drive pressures.
Secondly, as the number of tubes, n increases, the drive pressure
required for a given boost of intraventricular pressure increases.
This is accompanied by a concomitant decrease in the fluid flow of
volume per stroke required of the hydraulic pump system. Third, the
hydraulic pressure required increases as the tubes are
progressively inflated from being fully deflated to fully
circular.
[0104] Illustrated in FIG. 7 is the ratio of the hydraulic to the
afterload pressure as a function of the parameter of tube inflation
angle 2 for n=10. In the curve of FIG. 7, the targeted operating
point is shown by the asterisk. As can be seen a significant
contraction is achieved when theta is between 1.2 to 1.4 radians,
representing a 22% to 30% contraction. Thus, the operational range
of the hydraulic pressure is near one atmosphere for full assist
against an afterload pressure of approximately 90 mm Hg, and 1/4
atmosphere for cardiac boosting, that is, increasing the
ventricular pressure by 20 to 30 mm Hg.
[0105] As discussed earlier the artificial myocardium requires
synchronization of its contraction with the natural heart.
Contraction of the device must be timed appropriately with the
heart's systole. Additionally, the drive pressures during systolic
ejection must be maintained to match the needs of physiologic
afterloads.
[0106] The first factor, that is, the timing, can be achieved by
implanting an epicardial lead in a myocardial region of the natural
heart not in contact with the artificial myocardium, which could be
near the apical region, or at the right atrial appendage. The
artificial myocardium would be timed to contract with the R wave
produced on this lead. This detection can employ hardware that is
used at the present time in implantable defibrillators. The
systolic duration (in milliseconds) is preprogrammed to match its
functional dependence to the beat rate (BR) in beats per minute
expressed as .tau..sub.s=549 msec-2.times.BR.
[0107] In FIG. 8 the timing relationship between the ECG and the
artificial myocardium is represented graphically. .tau..sub.R is
the delay time between the start of the artificial myocardium
systole relative to the R wave. .tau..sub.S is the artificial
myocardium systolic duration, and .tau..sub.d is its diastolic
duration. The exact coincidence of the start of the diastolic
duration and the ECG T-wave is not critical. Large deviations from
this coincidence could either provide insufficient support
(.tau..sub.d starts too early) or hamper diastolic filling
(.tau..sub.d starts too late). Of course, irregular rhythms in the
natural heart, such as the occurrence of bigeminy, premature
ventricular contractions (PVC), or transient arrhythmias can also
affect performance of the artificial myocardium system. The most
straightforward way of dealing with this situation is to cause the
artificial myocardium wrap to be immediately deflated to assume the
diastolic state when this occurs. It can also be arranged so that
when no ECG trace is detected, the device would contract at beat
rates consistent with maintaining physiologic filling pressures.
Thus in the extra-cardiac support system of this invention certain
unique operating characteristics can be provided when working in
conjunction with the natural heart. It can, for example, provide
extra contractility not present in other support approaches.
[0108] FIG. 9 illustrates the left ventricular pressure/volume
relationship. In FIG. 9 the left ventricular pressure is plotted
against the left ventricular volume. The solid curve (loop 1)
illustrates the performance of the natural heart, while the dotted
curves show how the pressure volume relationship may be altered in
the presence of the extra cardiac assist device. There are two
factors that change as a result of the support device: the systolic
pressure and the stroke volume. The optimal assist mechanism for
this device is to boost the systolic pressure while allowing the
myocardium to retain its isovolumetric characteristics by elevating
the epicardial pressure (P.sub.e) and increasing transmyocardial
pressure (P.sub.t). The net effect is to displace the isovolumetric
curve upwards by P.sub.e. This is illustrated by the dashed line A
shown in FIG. 9. Whether the intraventricular pressure (P.sub.V)
remains at P.sub.t, achieves the maximum value of P.sub.t+P.sub.e,
or more likely reaches somewhere in between the two extremes,
depends on the vascular resistance and the ventricular stroke
volume. FIG. 9 illustrates the three possibilities. In loop 2, the
elevated ventricular pressure is matched by an increase in the
afterload. This results in no change in stroke volume. In loop 3,
the afterload remains unchanged, while the stroke volume is
increased by the assist device. Case 2 would result if the
ventricular stroke volume is equal to or greater than that
available from the device. This represents, nominally, a healthy
heart which requires no assist. The control scheme will be based on
achieving a full systolic contraction even though the diastolic
ventricular filling may not be complete on a beat by beat
basis.
[0109] With this control scheme, for a healthy heart which can
generate sufficient tension against physiologic afterloads, the
contractility required of the artificial myocardium would be zero,
and minimal hydraulic power would be required to fill the tubes.
This in turn generates minimal epicardial pressures and the assist
is at a minimum. However, in cases where the natural heart is not
capable of generating normal ejection fractions, the device
contraction will extend the stroke volume of the ventricle. In
order to realize these higher stroke volumes from the ventricle,
the artificial myocardium must provide the additional contractility
needed. This requires higher hydraulic power resulting in higher
epicardial pressure translating to higher ventricular pressure.
Under such circumstances, assist will increase flow and the
afterload will also increase as a result of the increased flow.
This is illustrated by curve 4, the most likely operating P-V loop
under assist conditions.
[0110] The control scheme is relatively straightforward. The device
will be operating at a full systolic stroke in every beat. Power
required to achieve the full stroke will be adjusted on a beat by
beat basis. The hydraulic stroke volume will be measured on every
beat in order to permit implementation of this algorithm. During
diastole, hydraulic pressure will be measured to provide an
indication of the end diastolic pressures. This information will be
used to determine system beat rate for a heart with no rhythm.
[0111] FIG. 10 illustrates in block diagrammatic form the hydraulic
system including the artificial myocardium 11, the energy converter
19 and the hydraulic reservoir 21. In one embodiment, the
artificial myocardium 11 consists of four layers of
polyetherurethane (Angioflex.RTM.) reinforced by polyester mesh,
fabricated from 40 micrometer fibers. Two layers of the fiber are
interwoven to yield tubular interconnections. Separators are
inserted in the individual tubes to prevent the opposing inner
walls from bonding with each other during the manufacturing
process. The process yields a device that has an overall wall
thickness of approximately 1/2 mm and a strength of 150 lbs/in,
nearly two orders of magnitude higher than the tensile forces
experienced by the device. Attached to the epicardium of the
natural heart, the device, when deflated, represents no additional
diastolic resistance to the heart. The energy converter 19 shuttles
fluid between the flexible reservoir 21 and the artificial
myocardium 11. Fluid reversal is achieved by a rotating valve. The
system flow resistance is designed to be low such that in the event
of a stoppage of the system, fluid in the tubes of the artificial
myocardium will automatically empty into the reservoir 21 within a
few heart beats. The positive diastolic filling pressure of the
natural hem and the negative intrathoracic pressure insures a
driving force to empty the fluid from the device. Once it is
completely emptied the artificial myocardium becomes a highly
flexible sheet which follows the wall motion of the natural hem
without any additional resistance. Such a device can be restarted
without any fear of embolic complications after a temporary
stoppage.
[0112] The direction of hydraulic fluid flow during systole in this
cardiac assist device is from the compliant fluid reservoir,
through the distributing manifold, and into the individual tubules
of the artificial myocardium. During diastole these bladders must
be emptied by reversing the direction of hydraulic fluid flow and
pumping fluid back into the reservoir. The direction of fluid
movement will be reversed by a rotary porting valve in conjunction
with unidirectional operation of the centrifugal pump itself.
[0113] Unidirectional pump operation has several advantages over
reversing the pump direction. Principally, unidirectional rotation
of the pump shaft presents much more favorable conditions for
bearing life. Although unidirectional pump speed will most likely
change between systole (artificial muscle inflation) and diastole,
these accelerations and deceleration represents a fraction of those
that would be associated with complete reversal of pump direction.
Furthermore, in a unidirectional mode, pump impeller design can be
optimized for fluid motion in one direction.
[0114] Some of the key dimensions and motor performance parameters
for a suitable motor for this system are listed below:
[0115] overall diameter: OD=1.0 in
[0116] overall length: OL=0.5 in
[0117] torque constant: KT=0.4 oz-in/amp
[0118] voltage constant: K.sub.B=0.3 V/kRPM
[0119] terminal resistance: R.sub.M=0.325 ohms
[0120] viscous damping: FI=0.002 oz-in/IRPM
[0121] The motor performance parameters listed above were used to
predict the anticipated torque-speed and efficiency characteristics
of the motor. The torque-speed performance characteristics indicate
that the specific operating point of 1.2 oz-in at 35000 RPM can be
obtained with an applied voltage of 12 V. In this application, the
supply voltage will be larger than 12 V, therefore pulse width
modulation (PWM) of the motor supply can and will be used to
control speed.
[0122] Unidirectional pump operation implies that fluid flow
reversal will be accomplished using a rotary porting valve. As
illustrated in FIGS. 11A and 11B, a balanced inflow and outflow
porting configuration is designed to minimize radial loads. The
rotary porting valve 32 consists of two concentric sleeves 34a and
34b, a fixed inner sleeve (not shown) and an outer sleeve 34a which
can be rotated through a defined angle by a torque motor (not
shown). The valve is composed of two pairs of inlet ports leading
to the impeller intake, one pair 40 coming from the fluid reservoir
and the other pair 38 from the hydraulic cuff 11, and two
corresponding pairs of outlet ports off the impeller, one leading
to the distributing manifold and one to the fluid reservoir.
Accordingly, the valve is designed so that switching the outer
sleeve into the systolic position opens the inlet port from the
reservoir while closing that from the manifold, and opens the
outlet port to the manifold while closing that to the reservoir.
Conversely, switching the outer sleeve back reverses the inlet and
outlet ports and generates diastole. The rotary porting valve also
incorporates hydraulic dampers to prevent valve rebound at the end
of switching. FIGS. 11A and 11B illustrate this valving scheme.
[0123] The pump motor bearing is a component which requires careful
design considerations By the choice of a unidirectional pump, the
major failure mode of bearings due to pump reversals inherent in
some designs has been eliminated. However, the artificial
myocardium operates at a high RPM in the range of 5,000 to 35,000
RPM. This high RPM places a stringent requirement on the motor
bearings.
[0124] Bearing load reduction can be achieved by judicious design.
Although the outflow pressure is high in the AMA, the surface areas
of the energy converter housing exposed to the pressure difference
is relatively small, such that the load on the bearings generally
remains in the same range. The use of a symmetric paired porting
design for the inflow and outflow orifices, allows the radial loads
to be reduced to zero.
[0125] For any extra cardiac support device, a volume compensating
chamber for the actuating volume is required. Since both the left
and right sides are not completely independent of the other, the
extra cardiac support cannot alternatively pump the left and the
right side. In general, for extra cardiac support, the two sides
have to be pumped simultaneously. The artificial myocardium
inherently has a volume reduction factor requiring only 25 cc of
the hydraulic stroke volume for a biventricular support. This
compensating chamber can be directly incorporated on the energy
converter as a flexing member or it can be separated from the
energy converter via a hydraulic conduit and shaped as a pancake
flexible chamber. The latter is the preferred approach since this
provides additional flexibility in chamber placement. This chamber
would be a 21/2" diameter sac with flexible membranes to yield a 1
cm excursion. The body of the chamber can be made from
Angioflex.RTM.. A velour layer can be the enclosure to provide
tissue layer stability.
[0126] The control of the proposed artificial device needs to
incorporate three key modes. These are, (1) synchronized
contraction and dilatation with the heart when a normal R-wave is
discerned, (2) no pumping during intermittent arrhythmia, and (3)
pumping at rates determined only by filling pressures during
fibrillation and cardiac arrest. An additional design criterion is
to ensure that the device does not work against the heart. The
proposed control algorithm consists of three modes discussed
below.
[0127] Synchronization is achieved by sensing the rhythm of the
natural heart, or paced signals for subjects with implantable
pacers. Two basic approaches are available, using either the P-wave
or the R-wave, the choice being governed by the conduction
capability of the heart.
[0128] P-wave may be preferable as the reference for
synchronization, since the right atrium remains free from any
mechanical contact with the device. Naturally, if a subject suffers
from frequent atrial flutter or atrial fibrillation, reliable
P-waves would not be available. In addition, AV block would exclude
the use of the P-wave. Patients with these pathologic conditions
would require R-wave sensing. Epicardial leads can be used for
either sensing mode. For atrial sensing, a lead can be sutured to
the atrial appendage, while ventricular sensing can be achieved by
a corkscrew electrode attached near the apex where no direct
squeezing of the myocardium would occur. Bipolar electrode designs
may be used in order to localize signal reception especially,
P-waves, with reduced noise pick-up in the acquired signals.
Unipolar leads can be used for R-wave sensing since this is the
simplest type of electrode for ventricular epicardial fixation.
[0129] Whether P-wave or R-wave sensing is used, the algorithm is
designed for synchronous contraction of the artificial device and
the myocardium. For a subject with a regular heart beat, this can
be achieved readily. For P-wave sensing, the device systole is
timed to initiate after .apprxeq.160 msec, a normal AV delay,
following P-wave detection. With R-wave detection, device actuation
is initiated immediately. An anticipation algorithm which is based
on the prior R-R intervals can also be used. Such algorithms are
available.
[0130] In FIG. 12A and 12B there is illustrated a subdermal port
for the artificial myocardium assist system in case there is a
failure of the hydraulic pumping capacity. FIG. 12A is a cross
sectional view across the skin interface 45 and FIG. 12B is an
"x-ray" view of the system viewed externally. In the case of a
system failure, as a result of electronic or mechanical problems,
the subdermal port can be accessed through a skin puncture with an
array 48 of 15 gauge needles. The procedure would involve the
extraction of the hydraulic fluid using a 50 cc syringe. This
extraction would collapse the artificial myocardium cuff 11. A hand
operated pneumatic pump (not shown) could then be connected to the
needle manifold 47 to activate the artificial myocardium. The
reason for the extraction of the hydraulic fluid and subsequent
manual use of a pneumatic pump is that the flow resistance through
a 1 cm long parallel array of 15 gauge 1 mm ID) needles is less
than 20 mm Hg for air, while the use of hydraulic fluid would
result in pressure losses which are orders of magnitude higher. The
artificial myocardium system would be implanted through a median
sternotomy. This procedure is sufficiently simple so that it would
be possible without bypasses, although severely compromised
patients might require bypass for support during the surgical
procedure. Of course, other perhaps less invasive, surgical
techniques could possibly be employed for this implantation. An
appropriately sized artificial myocardium cuff 11 would be wrapped
around the natural heart. The energy converter 19 and the hydraulic
reservoir 21 would be implanted in the thorax. It is estimated that
the total volume and weight of the thoracic unit would be
approximately 105 cc and 165 g respectively. The energy converter
and the fluid reservoir, which in practice could be an integral
part of the energy converter, would be anchored to the rib cage
with a flexible hydraulic connection to the artificial myocardium
cuff 11 and an electrical cable tunneled through the costal
diaphragmatic region to the electronic components which could be
implanted in the abdomen. These implant locations are illustrated
in FIG. 2. It is important that the artificial myocardium cuff 11
is anchored properly relative to the natural heart such that during
systolic contraction, the heart would not slip out of the
myocardium cuff 11. Suitable attachment arrangements are, for
example, illustrated in U.S. Pat. No. 4,957,477.
[0131] The primary biocompatibility issue for the artificial
myocardium relates to the epicardial tissue/cuff interface, or
pericardial tissue/cuff interface. The material in contact with the
epicardium or the pericardium will be the polyetherurethane
material Angioflex.RTM.. Other implanted components would consist
primarily of cable jackets made from either medical grade room
temperature vulcanizing rubber (RTV) or Angioflex.RTM.
polyurethane. Non-flexing parts would consist of titanium as casing
for the energy converter and electronics packages. Infectious risks
would be minimized in this design by the elimination of
percutaneous exit and entry sites into the body, by quality control
of surfaces and by choices of materials in contact with the
tissue.
[0132] FIGS. 13 and 14 show the calculated pressure and volume
ratios of the hydraulic and physiologic blood system fluids as a
function of the number of tubes in the artificial myocardium. FIG.
13 shows the hydraulic physiological pressure ratio as a function
of the number of tubes for both biventricular support and
univentricular support. FIG. 14 shows the physiological to
hydraulic volume ratio. These figures show that as the number of
tubular segments increases, the required hydraulic stroke volume
decreases while the required hydraulic pressure increases. For a
volume ratio of three, the univentricular support (1/2 of the total
wrap) would require 11 segments, while the biventricular support
would need 23 segments because of the larger perimeter for
biventricular support. This volume amplification between the
hydraulic stroke volume and the blood stroke volume is very
significant, since it permits the actuating system for the
artificial myocardium to be small and compact in size. In order to
take advantage of this volume amplification, the pressure required
to inflate the tubes to the appropriate extent for a significant
stroke work is approximately 10 times the afterload pressure of the
blood being pumped. In the figures, the intersections of the
univentricular and biventricular supports with the dashed
horizontal line indicate these operating points. FIG. 15
illustrates the blood stroke volume in cc's as a function of the
number of tube segments (n) in the artificial myocardium. As
illustrated, the stroke volume does not have a strong dependence on
n, especially when n becomes large and the hydraulic displacement
component becomes negligible compared with the contractile effect.
For a completely failed ventricle, in order to generate 100 mm Hg
of systolic pressure, the hydraulic drive pressure required will be
approximately 1,000 mm Hg. This value is illustrated by the star
shown in FIG. 7. The benefits derived from operating the energy
converter 19 at lower flow and higher pressures are higher system
efficiency as a result of lower flow losses and smaller system size
due to the lower volume requirement. The gain in the hydraulic
efficiency will be primarily in the energy converter 19. For
artificial myocardium cuff 11, the flow velocities in the
individual tubes is independent of the number of tubes. The volume
of each tube scales with the area of the tube so that the flow
velocity is a parameter determined only by physiologic
requirements. For a 65 cc stroke volume at a beat rate of 140, the
peak hydraulic flow velocity in the tubes is approximately 15 cm
per second, resulting in a dynamic pressure of approximately 0.1 mm
Hg, which has no impact on efficiency when compared to driving
pressures on the order of one atmosphere. With this design, the
wall stresses in the artificial myocardium cuff 11 are, as in flow
losses, independent of the number of tubes used in the wrap. The
wall stresses in the walls connecting the tubes are only functions
of the physiological parameters, such as the heart diameter and the
physiological pressure. The wall stress is given as the product of
the hydraulic pressure PH and the tube radius, r, and is a
constant, independent of the number of tubes, n.
[0133] The design consideration for the number of tubes per wrap
will be determined by practical considerations such as the
fabrication techniques and energy converter efficiency.
[0134] FIG. 16 shows a mock loop used for in vitro tests of the
artificial myocardium. Fluid from the reservoir 72 enters an atrium
68 which empties through a inflow valve 70 to the ventricle 64. The
ventricle consists of a cylindrical bladder which is surrounded by
another concentric cylindrical pouch. A space between the bladder
and the pouch is filled with a viscous fluid to simulate the
ventricular wall. The exterior of the pouch has fitted eyelets
spaced to accept an artificial myocardium simulating a left
ventricular wrap. The outflow from the artificial ventricle 64 is
coupled through another tri-leaflet valve 82 to an aortic
compliance chamber 80, followed by a flow probe 78 and a flow
rotameter 76. The outflow resistance 74 is adjustable. The return
flow empties into the reservoir 72. For this study the artificial
myocardium assist system was actuated using a pneumatic drive
console consisting of a high pressure plenum and a low pressure
plenum which were alternately switched to the device by solenoid
valves initiating systole and diastole respectively. This drive
mechanism replaces the hydraulic energy converter which would be
employed in the implantable system. For this study FIG. 17
illustrates a linear relationship between the afterload pressure
(AOP) and driving pressures for the artificial myocardium. The
theoretical calculated value is shown as the solid curve, while the
square dots indicate the values determined in this experimental
study. The flow output was maintained constant by adjusting the
aortic resistance. In the illustrated set of measurements the flow
was maintained at 6.5 liters per minute, with a filling pressure of
14.6 mm of mercury, a beat rate of 169 beats per minute and a
systolic duration and duty factor of 168 milliseconds and 40%
respectively. The device used in this study had 7 adjacent tubes
and provided 50% wrapping of the pouch simulating the natural heart
ventricle. The diameter and length of the pouch were 6 cm and 5 cm
respectively, fairly typical of a small, left ventricle. Based on
calculations, the anticipated stroke volume was 46.7 cc and the
measured stroke volume was 38 cc, 82% of the theoretical value.
[0135] A second set of measurements was obtained in this study by
maintaining a constant afterload, while the drive pressure was
varied and the resultant ventricular flow recorded. The results of
this study are illustrated in FIG. 18. The solid curve shows the
calculated values and the set of square points illustrate the
measured values. The conditions were similar to those for the
experiment illustrated in FIG. 17. FIG. 18 also illustrates the
calculated flow versus the drive pressure relationship. The outflow
pressure was set at 115 mm Hg, which is the intercept of the drive
pressure at zero flow. This study showed that the experimental
pneumatic drive pressure was slightly higher than that which would
have been predicted by the theoretical calculation. The data from
this study indicates by controlling the drive pressure, which is
equivalent to adjusting the contractility of the artificial
myocardium, both flow and pressures can be enhanced.
[0136] In FIGS. 19A and 19B there is illustrated one embodiment of
a girdle for wrapping around a heart to constrain dilatation of the
ventricle and limit the amount of energy and oxygen required to
maintain the heart muscle in tension.
[0137] In FIG. 19A the natural heart 110 is shown with the left
ventricle 112 somewhat dilated and with a girdle 117 surrounding
both the left ventricle 112 and the right ventricle 111. The girdle
117 is formed as illustrated in FIG. 20, with a series of
horizontal segments 113a-113i encircling the heart 110, the
segments toward the apex of the heart being smaller in cross
section and in length. The girdle segments 113 are filled with
hydraulic fluid which is maintained at a constant volume during the
beating of the heart. In this arrangement the girdle is entirely
passive and a distensible girdle lining 118 conforms to the shape
of the heart at the myocardium girdle lining interface by virtue of
the pressure of the fluid filled segments 113 against the
distensible inner lining 118. As shown in FIG. 3, when this girdle
is implanted around a natural heart the volume is controlled
through a three-way valve 127 which controls the amount of fluid
supplied to the girdle segments 113 from reservoir 125, which is
formed of a rigid casing 124.
[0138] According to equation (1), it can be seen that an increase
in mechanical work by a large factor results in a small increase in
oxygen consumption, but an increase in tension time causes a large
increase in oxygen consumption. Passive girdling of the heart, as
illustrated in FIGS. 19-21, acts to limit or reduce the ventricular
size of the diseased ventricle. Over an extended period of time,
which may be days or weeks, the fluid 114 volume may be increased,
thereby decreasing the periphery of the interface lining 118 of the
girdle, which may over a period of time actually decrease the
dilatation of the ventricle 112.
[0139] In FIG. 21 a control system for controlling the fluid
pressure in the segments 113 according to the tension in liner 118
is shown. The fluid pressure in girdle 113 is controlled by a
feedback loop including a strain gauge 142 placed at the interface
between the inner lining 118 and the myocardium providing a sensed
value for the tension of the myocardium, to hydraulic actuating
electronics 122 which may be a conventional hydraulic control
circuit. The electronic actuator 122 controls a conventional
mechanical fluid actuator 123 which provides for increase or
decrease of fluid within the girdle 117. This actuator operates in
conjunction with a three-way valve 127 and fluid reservoir 125. The
change in volume effected by this feedback, is not intended to, nor
does it operate in the time frame of the beating of the natural
heart. It is meant to adjust the volume over a much longer time
period, typically days, weeks or months.
[0140] In this configuration, the series of generally cylindrical
segments 113 are typically formed of non-distensible material. They
are attached to one another along the long axis of the cylinder and
may be filled with fluid either individually or in parallel. When
the fluid volume within the compartments 113 is very low, then the
girdle 113 assumes the shape shown in FIG. 19A providing for a
large inner diameter. On the other hand, when the fluid volume is
increased the segments assume, at fill inflation, a circular cross
section thereby decreasing the inner perimeter very substantially,
as illustrated in FIG. 19B. Thus, by controlling the volume of the
fluids supplied to the individual segments 113, the inner diameter
of the girdle 117 can be adjusted to be a close fit to the natural
heart. This configuration has the advantage that, since there is no
single vertical compartment, there is no gravity pooling of fluid
in one portion of the girdle 117. FIG. 22 illustrates a second
embodiment of this invention. The girdle 130 of FIG. 22 is an
adjustable girdle made from a synthetic material that can limit
tension, but is otherwise deformable to conform to the anatomical
geometry of the heart. In this case, the girdle 130 is formed of a
confining net 132 which is wrapped around the heart from the apex
to the atrioventricular (A-V) groove. The purpose of this net is to
limit the maximum diastolic dimension of the heart, while offering
no resistance to systolic ejection. In the design illustrated in
FIG. 22 a number of interlinked two-dimensional loops such as
lightweight plastic rings 133 are interconnected to form the girdle
or wrap 130. The loops 133 are free to move in all directions
without restraint, since none are physically connected to each
other. Rather, they are interlocked by having the loops or rings
133 pass through one another. The design of FIG. 22 presents no
systolic load to the contracting heart. The loop-mesh 132 can
readily conform to the shape of the heart with the change in
surface area accompanying the heart contraction readily
accommodated by the free loops.
[0141] An alternative form of this loop-mesh girdle is shown in
FIG. 23. In FIG. 23 a string system 134 is included with the string
attached to the loops 133 to effect change in the size of the mesh
by virtue of pulling the strings. This arrangement is able to
accommodate a treatment modality for scheduled size reduction to
the heart over a suitable period of time. In FIG. 23, a segment of
the girdle or wrap 130 is shown. The original size of the wrap can
be seen at the wide edge 136, while the narrowed down section is
seen at the ridge 138 of the wrap. Pulling on the two ends of two
sets of strings reduces the size of the mesh in two directions.
This can be done during a thoracoscopy or through a cutaneous
access port. In the construction illustrated in FIGS. 22 and 23 the
net 130 will be attached at several attachment points, typically 4
to 6 in number, at the A-V groove and also perhaps near the apex of
the natural heart. At the original implant the surgeon will
optimize the fit to the heart as it is existing and will adjust the
size through the mechanism described above. This design will
accommodate spontaneous heart size reduction even though some parts
of the mesh may adhere to the epicardium. However, due to relative
motion between the loops, it is unlikely that the mesh will become
fully encapsulated. In FIG. 24 there is shown a girdle in
accordance with this invention which is formed of a sheet of an
expanded polytetrafluroethylene (PTFE) material 124, prestressed
such that it remains below its elastic limit and its tension in the
plane of the sheet is sufficient to create radially inward forces,
thus resisting expansion while permitting inward compression. In
other words the girdle will resist further expansion while
fittingly accommodating shrinkage. Other materials may be employed,
provided that they exhibit the above elasticity
characteristics.
[0142] In FIG. 25 there is illustrated a cross sectional view of a
tissue engineered girdle lining having a polymer scaffold 131 which
has been seeded with myocardial cells harvested from the recipient
mounted on a polymer substrate 131, the substrate either facing a
girdle structure or forming the inner surface of that girdle. The
tissue engineered lining faces the patient's myocardium. Such a
lining reduces the irritation which may occur between the
epicardium and artificial materials employed to form the girdle
itself. The lining 130 would, over time, integrate biologically to
the patient's myocardium.
[0143] Techniques for cell scaffold engineering are described in
the literature. Two examples being, Biodegradable Polymer Scaffolds
for Tissue Engineering by Lisa E. Freed, Gordana Vunjak-Novakovic,
Robert J. Biron, Dana B. Eagles, Daniel C. Lesnoy, Sandra K. Barlow
and Robert Langer and Tissue by Robert Langer and Joseph P.
Vacanti, Biotechnology, Vol. 12, July 1994 and Tissue Engineering,
Robert Langer and Joseph P. Vacanti, Science, Vol. 260 May 14,
1993.
[0144] This tissue engineering techniques may also be employed with
respect to other artificial materials which come in contact with
the heart in various surgical situations including the active
devices described hereinabove and in U.S. patent application Ser.
No. 08/490,080, filed Jun. 13, 1995.
[0145] Having described the above specific embodiments of this
invention, other embodiments implementing the concepts of this
invention will doubtless occur. While specific details of an
artificial myocardium and artificial myocardium assist system have
been illustrated, it will be understood that other embodiments may
be formed employing the principles of this invention.
* * * * *