Cardiac assist device

A muscle-powered pump to assist the natural heart is disclosed. The device comprises an oblate, spheroidal-shaped pumping chamber surrounded by innervated muscular tissue. The device may be coupled to the ventricle and descending aorta with valves and be stimulated in synchrony with the natural depolarization of the heart or the device may be inserted into the descending aorta and used as a counter pulsation device. In this application, the innervated muscle is stimulated after a brief delay from the natural cardiac depolarization.

BACKGROUND OF THE INVENTION 
The present invention relates to a totally implantable muscle-powered 
cardiac assist device to be used as an auxiliary pump in conjunction with 
the natural heart. In one configuration the device comprises a pair of 
tubular shunts coupled to the aorta and left ventricle of the heart which 
communicate with an elastic chamber formed in the shape of an oblate 
ellipsoid. Valves located within the shunts permit blood to flow from the 
weak or damaged left ventricle of the heart into the aorta when the 
elastic chamber is compressed. An alternate configuration involves the use 
of the elastic chamber as an extra-aortic counterpulsation device with no 
valve requirement. 
The mechanical energy required to compress the chamber is supplied by an 
innervated autogenous muscle surrounding the elastic chamber. This muscle 
is stimulated by an implantable pulse generator in synchrony with the 
ventricular depolarization of the patient's heart. In operation, the 
contraction of the elastic chamber under the influence of a muscle tissue 
forces blood into the aorta. Additionally, the pulse generator provides 
chronic ultra-low frequency stimulation to the muscle tissue to maintain a 
high population of slow twitch-type muscle fibers. 
The use of autogenous muscle to drive mechanical pumps is known in the art 
from U.S. Pat. No. 4,078,267 which discloses an artificial heart propelled 
by respiratory muscles. Devices of this type have enjoyed only limited 
success because, mammalian skeletal muscle is not capable of long-term 
pumping due to metabolic fatigue. Recently it has been demonstrated that 
chronic electrical stimulation of muscle tissue produces an adaptive 
transformation of muscle tissue which increases the capillary densitiy in 
the muscle tissue as well as the mytochondrial volume and results in an 
increased work capacity of the transformed muscle. Histologically, such 
tissue is transformed to the slow twitch-type which exhibits greatly 
increased resistance to fatigue. 
Early experimental evaluation of skeletal muscles for myocardial 
augmentation was reported by Kantrowitz and McKennon. See Experimental Use 
of the Diaphragm as an Auxiliary Myocardium, Surgical Forum 9, Page 266, 
1959. By wrapping diaphragm muscle around the heart and stimulating it via 
the phrenic nerve, they observed no significant hemodynamic effects; 
however, when employed as the counterpulsation device, they noted a 
short-term increase in the diastolic aortic pressure. Later, in 1964 
Nakamora and Glenn utilized the diaphragm to assist atrial function. The 
diaphragm graft in the atrium continued to contract in response to 
stimulation from the phrenic nerve and served to elevate the right atrial 
pressure chronically. See Graft of the Diaphragm as a Functioning 
Substitute for the Myocardium; an Experimental Study, J Surg Res 4; 435, 
1964. 
Other approaches which involve the use of small spring-loaded diaphragm 
pumps with externally positioned flap valves have been energized by canine 
quadricept femorous muscles. Mechanical pumps of this type have shown 
outputs of 600-700 milliliters per minute. 
These early studies demonstrated the potential for the use of skeletal 
muscles to augment ventricular action of the heart. However, this initial 
work indicated that a critical problem existed in the deterioration of 
muscle performance with continued use. Attempts at improving the 
hemodynamic behavior of the muscle graft by lower frequency stimulation 
was demonstrated by Doer, et al in 1984. See Synchronously Stimulated 
Skeletal Muscle Graft for Myocardial Repair, J Thorac Cardiovasc Surg 87: 
325, 1984. These more recent studies demonstrated that skeletal muscle, 
while initially capable of hemodynamic work, fatigues rapidly even under 
conditions less demanding than those which are tolerated indefinitely by 
the cardiac muscle itself. 
Although skeletal muscles contain populations of fibers which share many of 
the characteristics of cardiac muscle tissue, the skeletal type (I) or 
slow twitch fibers serve primarily a postural role in that they are 
required to sustain prolonged periods of activity without appreciable 
fatigue. However, in the tissue suitable for application to cardiac assist 
devices, these fibers are interspersed with at least an equal number of 
fast or type (II) fibers. These latter fibers have the properties suited 
to brief periods of intense activity, their fast contractile 
characteristics derive from specific contractile protein isoforms and 
extensive sacrotubular system and their dependence on energy derived from 
anaerobic glycolysis. This metabolic substrate renders the muscles 
susceptible to fatigue under conditions of prolonged use even at low 
cardiac rate duty cycles such as those demonstrated by Doer. Additionally, 
unlike cardiac muscle cells which contract as a synctyium, skeletal muscle 
fibers are normally recruited to an extent determined by the intensity of 
activation and in a fixed sequence. In practice, the fast fibers are the 
first to contract and the slow fibers are the last to contract. This 
structural property of the skeletal muscles minimizes the functional 
demand placed upon the fibers which are most susceptible to fatigue. 
However, the application of such tissues to cardiac assist devices require 
that all of the muscle tissues be recruited simultaneously and be equally 
active with the consequence of chronic fatigue. 
Over the past fifteen years, however, a plasticity of muscle fiber type has 
been demonstrated in response to chronic electrical stimulation. In 1969, 
Salmons, et al demonstrated that the contractile speed of fast muscles 
could be modulated to a striking extent by continuous electrical 
stimulation of the motor nerve at a frequency of 10 Hz. 
There is now a large body of evidence to show that fast skeletal muscles 
can ultimately acquire all of the physiological, biochemical, and 
ultrastructural characteristics of slow muscle under conditions of chronic 
stimulation. Such adapted muscles demonstrate a corresponding increase in 
the use of enzymes for aerobic metabolism and a decrease in the enzymes 
for glycolysis. 
When a change is also involved, the contractile proteins period is 
reflected by an increased conversion of light to heavy chain insoforms of 
myosin characteristic of slow muscle tissue. As these changes progress 
over a period of months, the muscle mass contracts progressively more 
slowly and is more resistant to fatigue than initially. These recent 
developments have suggested that appropriately adapted skeletal muscle may 
be harvested to restore myocardial function through surgical procedures. 
In the present application, however, chronically stimulated and transformed 
muscle tissue is utilized to actuate a biological pump implanted within 
the body and connected to the aorta for assisting a weakened or diseased 
ventricle in the delivery of blood to body tissues. At the present time, 
it is contemplated that the latissimus dorsi muscle will be dissected from 
the patient's back through a posterial aterial incision and mobilized, 
preserving its vascular and nervous structure. This pedicle will be passed 
into the thorax through a window created by the resection of approximately 
a 3 cm segment of the lateral arc of the second rib, thus permitting the 
insertiion of the pedicle into the thorax. The posterial lateral incision 
is closed, and access to the pedicle is achieved through a median 
sternotomy. The muscle flap will then be stretched along its striated side 
and wrapped around the elastic pumping chamber and closed at both ends 
using superficial interrupted sutures. After stretching the muscle flap 
around the elastomeric balloon, electrodes are then expected to be 
implanted, either on the nerve of the latissimus dorsi or through the 
muscle or both. The electrode system would then be attached to a suitable 
cardiomystimulator. 
In this context, the present invention is directed to an optimized 
biological pump which exploits the ability of transformed tissue to 
augment the ventricular action of the heart. This invention discloses two 
alternate embodiments to achieve the desired goal of a totally 
implantable, body-compatible cardiac assist system.

DETAILED DESCRIPTION OF THE INVENTION 
I. The first embodiment is referred to as a Apico-Aortic Conduit System 
(AACS), depicted in FIG. 1 at 10; and, the second embodiment is referred 
to as an Extra-Aortic Balloon Counterpulsation System (EABC) and is shown 
in the FIGS. 2, 3 and 4. 
In either embodiment, the pump consists of an elastomeric chamber 20, 
surrounded by a muscle sheath 22, formed from transformed muscle tissue. 
The chamber is shaped in the form of an oblate ellipsoid having a 
horizontal axis 26 and a vertical axis 24. In the AACS system, 
unidirectional heart valves 28, 30 may be provided to establish the flow 
direction of blood through the chamber. These values are located in 
apertures formed in the periphery of the elastomeric chamber. Valves 
suitable for this application include the Medtronic mitral heart valve 
Model 7700 having an orifice diameter of 2 cm for the entry valve 28. A 
valve suitable for the exit valve 30 of the chamber is the Medtroniac 
aortic heart valve Model A7700 having an orifice diameter of 1.6 cm. 
The elastic chamber is shaped in the form of an ellipsoid of revolution. 
The generating ellipse has a major or horizontal axis 26, which is the 
axis of revolution and a minor or vertical axis 24 as shown in the Figure. 
For a desired fluid stroke volume of 70 cc, the chamber should have a 
volume of approximately 140 cc. This is based upon an assumed ejection 
ratio of 50%. For a volume of 140 cc, the dimensions of the major and 
minor axes are related by b=5.78/.sqroot.a. 
To compute the minimum force required to pump the desired stroke volume, 
one may model the chamber as an equivalent cylinder, having a volume equal 
to the chamber, whose length is equal to the horizontal axis of the 
ellipse. In this instance, the cylinder will have a base radius b given by 
b=4.72/.sqroot.a. The force required to displace the desired blood volume 
is given by: (70/.tau.b).sup.2 (.rho./.pi.)=602.95a where .tau. is the 
ejection time (0.35 sec) and .rho. is the specific gravity of blood 
(1.055). This force corresponds to the end pressure or terminal pressure, 
P.sub.ter, in the chamber distributed over the exit aperture of the 
chamber as determined by the size of the aortic valve aperture, r.sub.o. 
In practice, sufficient muscle mass is wrapped around the balloon to 
generate a static pressure of 120 mm of mercury or 1.6.times.10.sup.5 
dynes per square, centimeter within the chamber. This is the available 
pressure, P.sub.av1, responsible for driving blood into the body systemic 
vessels. 
The mass flow rate for a Newtonian fluid in the laminar regime is given by 
Poiseuille expression 
##EQU1## 
where P is the pressure, L is the length of the tube, r is the tube 
radius, and .eta. is the viscosity coefficient. As previously mentioned, 
the available pressure responsible for driving the fluid out of the pump 
is related to the radius of the aortic valve as indicated by the 
relationship above. Likewise, the minimum pressure or terminal pressure in 
the chamber is related to the average radius of the balloon which is taken 
as the radius of the equivalent cylinder. The quantity Pr.sup.4 in the 
Poiseuille relation gives an estimate of the system compliance, and 
therefore, to achieve maximum compliance matching, we should have 
P.sub.av1 .times.r.sup.4 =P.sub.ter .times.(b).sup.4. This leads to: a 
r.sub.o.sup.6 =0.595. For an aortic valve orifice, r.sub.o, of 0.8 cm, we 
have a=2.27 cm and b=3.84 cm for the desired dimensions of the oblate 
ellipsoid. 
Optimization of the chamber size is based on a fluid flow rate, f, 
expressed in cc's per second, which is equal to the systolic's cardiac 
output of the cardiac assist device. The parameters should be optimized to 
provide a stroke volume of 70 cc, an ejection time of 350 ms and a volume 
flow rate of 200 cc per second. The fluid velocity is given by the flow 
rate divided by the cross-sectional area, A. Therefore, the average flow 
velocity during systolic time, v=f/A=2.86a. At the end of the ejection 
time, the fluid flow velocity within the chamber must become zero. 
With respect to the muscle mass 22 required in this cardiac assist device, 
one can use Young-LaPlace equation to compute the tension required at the 
wall of the chamber to generate the 120 mm of mercury pressure. For a 
cylindrical balloon of unit radius, the wall tension is computed to be 
1.6.times.10.sup.5 dyne per centimeter. Measurements of muscle fibers 
reveal that the isometric force generated by a tensed muscle is 
approximately 2.9.times.10.sup.3 grams per square centimeter of muscle 
cross-section or 2.9.times.10.sup.6 dynes per square centimeter of muscle 
cross section. See Casey, E. J.; "Biophysics, Concepts and Mechanisms," 
Reinhold Books, New York, 1962, p 262-294. Calculations for a cylindrical 
balloon of unit radius (R=1 cm) and sufficient length to accommodate at 
least 70 ml of blood leads to the following two useful rules of thumb: 
Rule 1: 
EQU M=2.times.10.sup.31 3 RP 
Where M is the muscular mass in grams, R is the balloon or bladder radius 
in cm, and P is the balloon pressure in dynes/cm.sup.2. 
Rule 2: 
EQU r.sub.o.sup.6 R.sup.-2 P=4.times.10.sup.3 
where r.sub.o is the radius of the tube connecting the balloon to the 
aorta. 
For example, in order to achieve the human systolic pressure of 120 mm Hg 
(1.6.times.10.sup.5 dynes/cm.sup.2) in a 70 ml balloon of radius R=1 cm, 
the required muscle mass is about 320 gm (11.3 oz.) according to Rule 1. 
Also, the radius, r.sub.o of the aortic valve, is estimated as 0.5 cm (0.2 
inch from Rule 2). 
By imposing an r.sub.o value of 0.8 cm and ejection ratio of 50% on the 
design parameters, it can be shown that the muscle mass required to wrap 
around the two caps of the oblate ellipsoid of volume 140 cm.sup.3 is 
approximately twice that required for a volume of 70 cm.sup.3, i.e., 645 
gm (23 oz.). 
II--The Extra-Aortic Balloon Counterpulsation Pump (EABC) 
The pumping chamber here needs no entrance or exit valves as shown in FIG. 
2 at 14 and in FIG. 4 at 16. The EABC chamber is connected directly to the 
divided left subclavian artery distal to the thoracodorsal and 
thoracoacromial branches. A series (T-connection) 14 or parallel 
(U-connection) 12, 16 pump can be used. The balloon can either be wrapped 
by the rectus abdominus and latissimus dorsi pedicles, or placed deep to 
the pectoralis major. 
The powering muscle would be stimulated directly by two wire electrodes 32. 
The stimulator is triggered from the left ventricular electrocardiogram 
via lead 34 or from the arterial pressure tracing output. Unlike the AACS, 
in this embodiment the pump would be triggered at the end diastolic phase 
of the cardiac cycle. This allows increased muscle perfusion which occurs 
while the muscle is relaxed during systole. Thus, fatigue can be 
considerably minimized, not only by this operational mode, but also by 
using the optimal stimulation parameters and protocol as with the AACS. 
In addition, the hemodynamic requirements for the EABC device are minimal. 
There are no valve requirements and the balloon volume can be chosen 
commensurate with the severity of the situation. A balloon volume of 30 to 
70 cc is recommended with an optimum size of 50 cc. The only requirement 
is that the balloon shape be spherical or nearly spherical in order to 
avoid sharp edges and corners where blood may stagnate. 
The EABC system can be made to offset the primary or essential 
hypertension. This type of high blood pressure is caused by the 
progressive increase in construction of arteries and arterioles and their 
decreasing compliance, a phenomena which gradually increases with age. 
This is to be distinguished from malignant hypertension which arises from 
hormonal disturbances of the adrenal glands that sit atop of the kidneys 
or from malfunctioning of the baroreceptors of the carotid sinus which is 
in the back of the neck. 
By adjusting the pressure wave on the extra-aortic balloon, one can augment 
the systolic pressure by decreasing the diastolic pressure level. Notice 
that infants average 80/46 in blood pressure at birth which rises to 
100/60 during the first ten days, and levels up at 120/70 during 
adulthood. The following increase seems to be gradual reaching 135/80 in 
the fifties and 150/85 in the seventies. The borderlines of 160/95 are at 
best empirical in the sense that they represent a gradual process, and a 
50 year-old subject with 160/90 blood pressure is the equivalent of a 
healthy counterpart who was 135/80 in his fifties and would extrapolate to 
160/90 at 90 or 100 years. The invention disclosed herein involves the 
gradual augmentation of the cardiac output in such a way to compliantly 
meet this progressive imbalance--with no extra demand from the heart 
muscle itself. 
The pulse generator 36 of the present device must be adapted to provide 
chronic background stimulation to the innervated autogenous muscle tissue 
to provide for the maintenance of a high type two fiber population. 
To provide for optimization of the stimulation parameters for any given 
individual it is required that the pulse generator be capable of providing 
burst stimulation with a burst duration between 150 and 500 milliseconds, 
with a number of pulses in a burst being less than or equal to 20. The 
pacemaker should also be capable of providing stimulation pulses at a rate 
between 0 and 150 beats-per-minute with a pulse width duration of between 
150 and 500 microseconds. To provide for adjustable thresholds of the 
autogenous tissue, it is desirable to have an amplitude adjustable within 
the range of 0 to 15 volts with constant current output. The device must 
have an R-wave synchronous or triggered operating mode for stimulating the 
autogenous muscle in phase with the depolarization of the cardiac tissue 
for use in configuration 10. The delay from ventricular sense to stimulus 
should be variable between 20 and 500 milliseconds and be programmable by 
the attending physician. 
It may also be desirable to provide for stimulating the autogenous tissue 
at a rate proportional to the sinus rhythm of the patient.