U.S. patent application number 10/613949 was filed with the patent office on 2004-05-20 for optimized pulsatile-flow ventricular-assist device and total artificial heart.
Invention is credited to Grandia, Lonn, Korakianitis, Theodosios.
Application Number | 20040097782 10/613949 |
Document ID | / |
Family ID | 25191792 |
Filed Date | 2004-05-20 |
United States Patent
Application |
20040097782 |
Kind Code |
A1 |
Korakianitis, Theodosios ;
et al. |
May 20, 2004 |
Optimized pulsatile-flow ventricular-assist device and total
artificial heart
Abstract
A method of optimizing a mechanical cardiac pumping device
includes modeling the circulatory system of the patient who will
receive the mechanical cardiac pumping device and identifying an
operating condition of the native heart to which the device will
respond. The model is used to determine the required blood volume
to be ejected from the device and an initial estimate of the power
required to be provided to the mechanical cardiac pumping device is
provided in order to provide the required ejected blood volume. The
resultant ejected blood volume is evaluated with data obtained from
the model and the estimate of the power requirement is then
updated. The above steps are iteratively performed until the power
required to obtain the necessary ejected blood volume is
identified. Possible variations of power and pumping rate that
allow the mechanical cardiac pumping device to provide the required
volume are determined and the variation that best matches the
physiological constraints of the patient and minimizes the power
required by the mechanical cardiac pumping device is selected. The
steps are iteratively performed until the mechanical cardiac
pumping device is optimized to respond to each desired operating
condition of the native heart.
Inventors: |
Korakianitis, Theodosios;
(St. Louis, MO) ; Grandia, Lonn; (Brussels,
IL) |
Correspondence
Address: |
BLACKWELL SANDERS PEPER MARTIN LLP
720 OLIVE STREET
SUITE 2400
ST. LOUIS
MO
63101
US
|
Family ID: |
25191792 |
Appl. No.: |
10/613949 |
Filed: |
July 3, 2003 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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10613949 |
Jul 3, 2003 |
|
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09805523 |
Mar 13, 2001 |
|
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6632169 |
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Current U.S.
Class: |
600/16 |
Current CPC
Class: |
A61M 60/40 20210101;
A61M 60/268 20210101; A61M 2205/3303 20130101; A61M 60/258
20210101; A61M 60/892 20210101; A61M 60/148 20210101; A61M 2205/33
20130101; A61M 60/562 20210101; A61M 60/50 20210101; A61M 60/896
20210101; A61M 60/432 20210101; A61M 60/894 20210101 |
Class at
Publication: |
600/016 |
International
Class: |
A61N 001/362 |
Claims
What is claimed is:
1. A method of optimizing the power required by a mechanical
cardiac pumping device in steady-state operating condition, said
method comprising the steps of: a. modeling the physical system, or
at least a portion thereof, of the patient who will receive the
mechanical cardiac pumping device; b. identifying an operating
condition of the native heart of the patient who will receive the
mechanical cardiac pumping device to which the mechanical cardiac
pumping device will respond; c. using the model of the physical
system from step a, above, to determine the required blood volume
to be ejected from the mechanical cardiac pumping device; d.
providing an initial estimate of the instantaneous power as a
function of time across at least one period of the heartbeat
required to be provided to the mechanical cardiac pumping device in
order to provide the required ejected blood volume; e. evaluating
the resultant ejected blood volume with data obtained from the
model of the physical system; f. updating the estimate of the power
requirement; g. iteratively performing steps e and f, above, until
the power required to obtain the required ejected blood volume by
the combined operation of the native heart and the VAD is
identified; h. determining possible solutions to the instantaneous
power as a function of time that allows the mechanical cardiac
pumping device to provide the required ejected blood volume; i.
choosing the solution from step h, above, that best matches the
physiological constraints of the patient and provides for optimal
power usage by the mechanical cardiac pumping device; and j.
iteratively performing steps b through i, above, until the
mechanical cardiac pumping device is optimized to respond to each
desired operating condition of the native heart.
2. A method of optimizing the power and energy required by a
mechanical cardiac pumping device in steady-state operating
condition, said method comprising the steps of: a. modeling the
physical system, or at least a portion thereof, of the patient who
will receive the mechanical cardiac pumping device; b. identifying
an operating condition of the native heart of the patient who will
receive the mechanical cardiac pumping device to which the
mechanical cardiac pumping device will respond; c. using the model
of the physical system from step a, above, to determine the
required blood volume to be ejected from the mechanical cardiac
pumping device; d. providing an initial estimate of the
instantaneous power as a function of time across at least one
period of the heartbeat required to be provided to the mechanical
cardiac pumping device in order to provide the required ejected
blood volume; e. evaluating the resultant ejected blood volume with
data obtained from the model of the physical system; f. updating
the estimate of the power requirement; g. iteratively performing
steps e and f, above, until the power required to obtain the
required ejected blood volume by the combined operation of the
native heart and the VAD is identified; h. determining the possible
solutions to the instantaneous power as a function of time, and the
total energy over the pumping cycle that allows the mechanical
cardiac pumping device to provide the required ejected blood
volume; i. choosing the solution from step h, above, that best
matches the physiological constraints of the patient and provides
for optimal power and energy usage by the mechanical cardiac
pumping device; and j. iteratively performing steps b through i,
above, until the mechanical cardiac pumping device is optimized to
respond to each desired operating condition of the native
heart.
3. A method of optimizing the energy required by a mechanical
cardiac pumping device in steady-state operating condition, said
method comprising the steps of: a. modeling the physical system, or
at least a portion thereof, of the patient who will receive the
mechanical cardiac pumping device; b. identifying an operating
condition of the native heart of the patient who will receive the
mechanical cardiac pumping device to which the mechnical cardiac
pumping device will respond; c. using the model of the physical
system from step a, above, to determine the required blood volume
to be ejected from the mechanical cardiac pumping device; d.
providing an initial estimate of the instantaneous power as a
function of time across at least one period of the heartbeat
required to be provided to the mechanical cardiac pumping device in
order to provide the required ejected blood volume; e. evaluating
the resultant ejected blood volume with data obtained from the
model of the physical system; f. updating the estimate of the power
requirement; g. iteratively performing steps e and f, above, until
the power required to obtain the required ejected blood volume by
the combined operation of the native heart and the VAD is
identified; h. determining possible solutions to the total energy
over the pumping cycle that allows the mechanical cardiac pumping
device to provide the required ejected blood volume; i. choosing
the solution from step h, above, that best matches the
physiological constraints of the patient and provides for optimal
energy usage by the mechanical cardiac pumping device; and j.
iteratively performing steps b through i, above, until the
mechanical cardiac pumping device is optimized to respond to each
desired operating condition of the native heart.
4. A method of optimizing the power required by a mechanical
cardiac pumping device in steady-state operating condition, said
method comprising the steps of: a. modeling the physical system, or
at least a portion thereof, of the patient who will receive the
mechanical cardiac pumping device; b. providing an initial estimate
of the instantaneous power as a function of time across at least
one period of the heartbeat required to be provided to the
mechanical cardiac pumping device in order to provide the required
ejected blood volume; c. evaluating the resultant ejected blood
volume; d. updating the estimate of the power requirement; e.
iteratively performing steps c and d, above, until the power
required to obtain the required ejected blood volume by the
combined operation of the native heart and the VAD is identified;
f. determining the possible solutions to the instantaneous power as
a function of time that allows the mechanical cardiac pumping
device to provide the required ejected blood volume; g. choosing
the solution from step f, above, that best matches the
physiological constraints of the patient and provides for optimal
power usage by the mechanical cardiac pumping device; and h.
iteratively performing steps b through g, above, until the
mechanical cardiac pumping device is optimized to respond to each
desired operating condition of the native heart.
5. A method of optimizing the energy required by a mechanical
cardiac pumping device in steady-state operating condition, said
method comprising the steps of: a. modeling the physical system, or
at least a portion thereof, of the patient who will receive the
mechanical cardiac pumping device; b. providing an initial estimate
of the instantaneous power as a function of time across at least
one period of the heartbeat required to be provided to the
mechanical cardiac pumping device in order to provide the required
ejected blood volume; c. evaluating the resultant ejected blood
volume; d. updating the estimate of the power requirement; e.
iteratively performing steps c and d, above, until the power
required to obtain the required ejected blood volume by the
combined operation of the native heart and the VAD is identified;
f. determining the possible solutions to the total energy over the
pumping cycle that allows the mechanical cardiac pumping device to
provide the required ejected blood volume; g. choosing the solution
from step f, above, that best matches the physiological constraints
of the patient and provides for optimal energy usage by the
mechanical cardiac pumping device; and h. iteratively performing
steps b through g, above, until the mechanical cardiac pumping
device is optimized to respond to each desired operating condition
of the native heart.
6. A method of optimizing the power and energy required by a
mechanical cardiac pumping device in steady-state operating
condition, said method comprising the steps of: a. modeling the
physical system, or at least a portion thereof, of the patient who
will receive the mechanical cardiac pumping device; b. providing an
initial estimate of the instantaneous power as a function of time
across at least one period of the heartbeat required to be provided
to the mechanical cardiac pumping device in order to provide the
required ejected blood volume; c. evaluating the resultant ejected
blood volume; d. updating the estimate of the power requirement; e.
iteratively performing steps c and d, above, until the power
required to obtain the required ejected blood volume by the
combined operation of the native heart and the VAD is identified;
f. determining the possible solutions to the instantaneous power as
a function of time and total energy over the pumping cycle that
allows the mechanical cardiac pumping device to provide the
required ejected blood volume; g. choosing the solution from step
f, above, that best matches the physiological constraints of the
patient and provides for optimal power and energy usage by the
mechanical cardiac pumping device; and h. iteratively performing
steps b through g, above, until the mechanical cardiac pumping
device is optimized to respond to each desired operating condition
of the native heart.
7. The method of any one of claims 1 through 6, wherein the modeled
physical system is used to determine the required blood volume to
be ejected from the mechanical cardiac pumping device by
discretizing the modeled physical system with Finite Element
Methods and Computational Fluid Dynamics into mass, damping, and
stiffness matrices, and their corresponding elemental
displacements.
8. The method of any one of claims 1 through 6, wherein the
physical system is modeled with MRI data and the modeled physical
system is used to determine the required ejected blood volume from
the mechanical cardiac pumping device by evaluating the MRI
data.
9. The method of any one of claims 1 through 6, wherein at least
some components of the physical system are modeled utilizing a
lumped-parameter model.
10. The method of any one of claims 1 through 6, wherein at least
some components of the physical system are modeled utilizing a
distributed-parameter model.
11. The method of any one of claims 1 through 6, wherein at least
some components of the physical system are modeled utilizing a
continuum model.
12. The method of any one of claims 1 through 6, wherein neural
networks are utilized to determine the instantaneous power as a
function of time, and the total energy over the pumping cycle, that
allows the cardiac device to provide the required ejected blood
volume.
13. The method of any one of claims 1 through 6, wherein heuristic
methods are utilized to determine the instantaneous power as a
function of time, and the total energy over the pumping cycle, that
allows the cardiac device to provide the required ejected blood
volume.
14. The method of any one of claims 1 through 6, wherein the
operating condition of the native heart to which the mechanical
cardiac pumping device will respond is heart rate.
15. The method of any one of claims 1 through 6, wherein the
operating condition of the native heart to which the mechanical
cardiac pumping device will respond is ventricular volume.
16. The method of any one of claims 1 through 6, wherein the
operating condition of the native heart to which the mechanical
cardiac pumping device will respond is ventricular pressure.
17. The method of any one of claims 1 through 6, wherein the
operating condition of the native heart to which the mechanical
cardiac pumping device will respond is at least a portion of the
ECG signal.
18. The method of any one of claims 1 through 6, wherein the
mechanical cardiac pumping device is a left ventricular-assist
device and the operating conditions of the native heart to which
the device will respond are at least one of: heart rate, heart
phase, left ventricular volume, right ventricular volume, left
ventricular pressure, and right ventricular pressure.
19. The method of any one of claims 1 through 6, wherein the
mechanical cardiac pumping device is a left ventricular-assist
device and the operating conditions of the native heart to which
the device will respond are at least one of: heart rate as a
function of time, heart phase, left ventricular volume as a
function of time, right ventricular volume as a function of time,
left ventricular pressure as a function of time, and right
ventricular pressure as a function of time.
20. The method of any one of claims 1 through 6, wherein the
mechanical cardiac pumping device is a right ventricular-assist
device and the operating conditions of the native heart to which
the device will respond are at least one of: heart rate, heart
phase, left ventricular volume, right ventricular volume, left
ventricular pressure, and right ventricular pressure.
21. The method of any one of claims 1 through 6, wherein the
mechanical cardiac pumping device is a right ventricular-assist
device and the operating conditions of the native heart to which
the device will respond are at least one of: heart rate as a
function of time, heart phase, left ventricular volume as a
function of time, right ventricular volume as a function of time,
left ventricular pressure as a function of time, and right
ventricular pressure as a function of time.
22. The method of any one of claims 1 through 6, wherein the
mechanical cardiac pumping device is a bi-ventricular-assist device
and the operating conditions of the native heart to which the
device will respond are at least one of: heart rate, heart phase,
left ventricular volume, right ventricular volume, left ventricular
pressure, and right ventricular pressure.
23. The method any one of claims 1 through 6, wherein the
mechanical cardiac pumping device is a bi-ventricular-assist device
and the operating conditions of the native heart to which the
device will respond are at least one of: heart rate as a function
of time, heart phase, left ventricular volume as a function of
time, right ventricular volume as a function of time, left
ventricular pressure as a function of time, and right ventricular
pressure as a function of time.
24. The method of any one of claims 1 through 6, wherein the
mechanical cardiac pumping device is a total artificial heart.
25. A method of optimizing the control scheme of a controller for a
mechanical cardiac pumping device, said method comprising the steps
of: a. simulating the steady-state physical condition of a patient
who will receive the mechanical cardiac pumping device; b.
identifying a new, target steady-state condition; c. determining
which outputs of the physical system to monitor to best perform the
transition from one steady-state to another; d. determining the
best combination of inputs, outputs, and modifications that achieve
transient transfer from one steady-state to another without
destabilizing the dynamic system; and e. storing the information
determined in steps c and d, above, in the controller.
26. A method of assisting the cardiac function of the native heart
of a patient using an implanted ventricular-assist device, said
method comprising the steps of: a. monitoring with a controller a
steady-state condition of the physical system of the patient having
the implanted ventricular-assist device; b. directing with the
controller the optimal power required to sustain the steady-state
while meeting physiological constraints; c. determining with the
controller when the physical system of the patient has left the
steady-state operating condition; d. determining with a controller
a new, target steady-state condition; e. while the physical system
of the is not in a steady-state operating condition, iteratively
performing the following steps i-v: i. monitoring inputs from the
physical system of the patient, the inputs being at least one of: a
measure of the heart phase, X/Xmax, and the shape of X/Xmax; ii.
evaluating with the controller the desired outputs from the
combined native heart and ventricular-assist device required at a
new steady-state condition, the outputs being at least one of:
heart rate, blood volume ejected by the native heart, blood volume
ejected by the ventricular-assist device, and the ECG trace; iii.
monitoring with a controller the actual outputs of the physical
dynamic system, the outputs being at least one of: heart rate,
blood volume ejected by the native heart, blood volume ejected by
the ventricular-assist device, and at least a portion of the ECG
trace; iv. modifying with a controller the actual output data
according to feedback transfer matrices stored within the
controller; v. transmitting with the controller modified inputs
from step iv, above, such that the desired outputs from step ii,
above, are achieved without destabilizing the dynamic system of the
patient during the transient period between steady-states; f.
iteratively performing the steps a-e, above, so long as the
ventricular-assist device is in operation.
27. A method of assisting the cardiac function of the native heart
of a patient using an implanted total artificial heart device, said
method comprising the steps of: a. monitoring with a controller a
steady-state condition of the physical system of the patient having
the implanted total artificial heart device; b. directing with the
controller the minimum power required to sustain the steady-state
while meeting physiological constraints; c. determining with the
controller when the physical system of the patient has left the
steady-state operating condition; d. determining with a controller
a new, target steady-state condition; e. while the physical system
of the is not in a steady-state operating condition, iteratively
performing the following steps i-v: i. monitoring inputs from the
physical system of the patient, the inputs being at least one of: a
measure of the heart phase, X/Xmax, and the shape of X/Xmax; ii.
evaluating with the controller the desired outputs from the
combined native heart and ventricular-assist device required at a
new steady-state condition, the outputs being at least one of:
blood volume ejected by the total artificial heart, and at least a
portion of the ECG trace; iii. monitoring with a controller the
actual outputs of the physical dynamic system, the outputs being at
least one of: blood volume ejected by the total artificial heart,
and at least a portion of the ECG trace; iv. modifying with a
controller the actual output data according to feedback transfer
matrices stored within the controller; v. transmitting with the
controller modified inputs from step iv, above, such that the
desired outputs from step ii, above, are achieved without
destabilizing the dynamic system of the patient during the
transient period between steady-states; f. iteratively performing
the steps a-e, above, so long as the total artificial heart device
is in operation.
28. A method of assisting the cardiac function of the native heart
of a patient using an implanted ventricular-assist device, said
method comprising the steps of: a. allowing the native heart to
pump as much blood as it is able prior to activation of the
ventricular-assist device; b. activating the ventricular-assist
device to provide additional pumping action as the blood-ejection
phase of the native heart nears completion such that the native
heart pumps more blood than it would unaided due to the reduction
of back pressure in the native ventricle caused by the pumping
action of the ventricular-assist device; c. coordinating the timing
of the action and length of the pumping stroke of the
ventricular-assist device with the ejected blood volume and rhythm
of the native heart such that the power required by the
ventricular-assist device is the optimal needed to pump the
required volume of blood while meeting physiological constraints;
d. varying the stroke displacement over time and resulting power
over time of the ventricular-assist device such that the power
required by the ventricular-assist device is the optimal needed to
pump the required volume of blood while meeting physiological
constraints; e. iteratively performing steps a through d, above, so
long as the ventricular-assist device is in operation.
29. A method of optimizing a mechanical cardiac pumping device
wherein unsteady fluid mechanics are used to optimize the forcing
function imposed by the mechanical cardiac pumping device such that
the power required by the mechanical cardiac pumping device is the
minimum power required to complement the cardiac output of the
diseased native heart, said method comprising the steps of: a.
modeling the dynamic response of the diseased native heart and of
the mechanical cardiac pumping device with experimental data; b.
using the instantaneous non-linear mass, [M], damping, [C], and
stiffness, [K] matrices of the dynamic model, and corresponding
elemental displacements {x} and its derivatives {x.} and {x}, as
inputs into an equation which sums these matrices to calculate the
forcing function, F{t}, of the dynamic system; c. calculating the
forcing function of the diseased native heart, F.sub.nh{t}; d.
calculating the required forcing function of the mechanical cardiac
pumping device, F.sub.vad{t}; e. inputing the value of
F.sub.vad{t}from step d, above, into a controller; and f.
connecting operatively the controller to a mechanical cardiac
pumping device, such that the controller is able to direct to the
mechanical cardiac pumping device the minimum power required to
achieve F.sub.vad{t}.
30. A method of optimizing a mechanical cardiac pumping device
wherein unsteady fluid mechanics are used to optimize the forcing
function imposed by the mechanical cardiac pumping device such that
the power required by the mechanical cardiac pumping device is the
minimum power required to complement the cardiac output of the
diseased native heart, said method comprising the steps of: a.
modeling the dynamic response of the diseased native heart and of
the mechanical cardiac pumping device with experimental data; b.
using the instantaneous non-linear mass, [M], damping, [C], and
stiffness, [K] matrices of the dynamic model, and corresponding
elemental displacements {x} and its derivatives {{umlaut over (x)}}
and {{dot over (x)}}, as inputs into an equation of the form:
[M]{{umlaut over (x)}}+[C]{{dot over (x)}}+[K]{x}=F{t}to calculate
the forcing function, F{t}, of the dynamic system; c. calculating
the forcing function of the diseased native heart, F.sub.nh{t}; d.
calculating the required forcing function of the mechanical cardiac
pumping device, F.sub.vad{t}, using an equation of the form:
F{t}=F.sub.nh{t}+F.sub.vad{t- }e. inputing the value of
F.sub.vad{t} from step d, above, into a controller; and f.
connecting operatively the controller to a mechanical cardiac
pumping device, such that the controller is able to direct to the
mechanical cardiac pumping device the optimal power required to
achieve F.sub.vad{t}.
31. A method of optimizing a mechanical cardiac pumping device
wherein unsteady fluid mechanics are used to optimize the forcing
function imposed by the mechanical cardiac pumping device such that
the power required by the mechanical cardiac pumping device is the
minimum power required to complement the cardiac output of the
diseased native heart, said method comprising the steps of: a.
modeling the dynamic response of the diseased native heart and of
the mechanical cardiac pumping device with experimental data; b.
using the instantaneous non-linear mass, [M], damping, [C], and
stiffness, [K] matrices of the dynamic model, and corresponding
elemental displacements {x} and its derivatives {{umlaut over (x)}}
and {{dot over (x)}}, as inputs into an equation of the form:
[M]{{umlaut over (x)}}+[C]{{dot over (x)}}+[K]{x}=F{t}to calculate
the forcing function, F{t}, of the dynamic system; c. calculating
the forcing function of the diseased native heart, F.sub.nh{t}; d.
calculating the required forcing function of the mechanical cardiac
pumping device, F.sub.vad{t}, using an equation of the form:
F{t}=F.sub.nh{t}+F.sub.vad{t- }e. balancing the instantaneous power
at any time t utilized by the mechanical cardiac pumping device
with an equation of the form: W(t)=F.sub.vad{t}.multidot.{{dot over
(x)}}+losses=V{t}.multidot.i{t}f. inputing the value W(t) from step
e, above, into a controller; and g. connecting operatively the
controller to a mechanical cardiac pumping device, such that the
controller is able to direct to the mechanical cardiac pumping
device the optimal power required to achieve F.sub.vad{t}.
32. A device to assist the function of a cardiac ventricle, the
device comprising: a. a first magnet having an open center and
formed of high ferromagnetic-constant material; b. a first vessel
surrounding the first magnet and defining a space in fluid
communication with the blood flow output great vessel of the
diseased ventricle of the patient using the device, the first
magnet being movable within the first vessel in substantially
fluid-tight relation thereto; c. a second magnet formed of high
ferromagnetic-constant material and in magnetic communication with
the first magnet so that the respective magnetic fluxes of the
first magnet and the second magnet affect each other, so that the
first magnet and the second magnet are biased toward and tend to
lock to one another, to thereby move in the same direction as one
another; d. a second vessel encasing the second magnet and defining
a space, the second magnet being movable within the space in
substantially fluid-tight relation to the second vessel, the space
defined by the second vessel being in fluid communication with a
hydraulic pump for actuating the second magnet; and e. an one-way
valve connected to the first magnet, the one-way valve being
movable with the first magnet, and adapted to cause movement of
blood from the diseased ventricle to and into the great vessel
associated with that diseased ventricle.
33. The device of claim 32, wherein the device is a L-VAD and is
sized and shaped for positioning between the aortic valve and the
aortic arch of the patient using the device.
34. The device of claim 32, wherein the device is a R-VAD and is
sized and shaped for positioning between the pulmonary valve and
the bifurcation of the pulmonary trunk of the patient using the
device.
35. A device (58, 74) to assist the function of a cardiac
ventricle, the device comprising: a. a first annular magnet (54)
formed of high ferromagnetic-constant material; b. a first sleeve
(68) surrounding the first annular magnet (54) and defining a space
in fluid communication with the blood flow output great vessel of
the patient using the device, the first annular magnet (54) being
longitudinally and reciprocally slideable within the first sleeve
in substantially fluid-tight relation thereto; c. a second annular
magnet (44) formed of high ferromagnetic-constant material and
sized and shaped for placement exterior of the first sleeve (68),
the second annular magnet (44) being disposed coaxially in relation
to and in magnetic communication with the first annular magnet
(54), so that the respective magnetic fluxes of the first magnet
and the second magnet affect each other, so that the first annular
magnet and the second annular magnet are biased toward and tend to
lock to one another, and to thereby move in the same direction as
one another; d. a second sleeve (72) encasing the second annular
magnet (44), the second annular magnet being longitudinally and
reciprocally slideable between the first sleeve (68) and the second
sleeve in substantially fluid-tight relation to the first sleeve
and the second sleeve, and the second sleeve (72) defining an
annular space (86) radially outwardly of the first sleeve (68) for
longitudinal travel therein of the second annular magnet (44), the
annular space (86) being in fluid communication with a hydraulic
pump for actuating the second annular magnet (44); and e. an
one-way valve (70) connected to the first annular magnet (54) and
disposed transversely in relation to the longitudinal axis of the
first annular magnet (54), the one-way valve being movable with the
first magnet, and closed when moving away from the origin of the
valve annulus when the device is in normal use position, to thereby
cause blood of the patient to move out of a diseased ventricle and
toward the great vessels associated with that diseased ventricle as
the first annular magnet (54) moves in a direction toward the great
vessels of the diseased ventricle due to magnetic flux of the
second annular magnet, the one-way valve further being adapted to
be open when moving in a direction away from the great vessels of
the diseased ventricle, to thereby permit blood of the patient to
flow through the one-way valve into the space defined by the first
sleeve (68) when the second annular magnet moves away from the
great vessels of the diseased ventricle.
36. The device of claim 35, wherein the entire device (58, 74) is
of sufficiently small size and weight for placement in normal
operative position between the valve annulus and at least a portion
of the great vessels of the diseased ventricle of a patient using
the device, to assist or replace the function of at least a portion
of the diseased native heart.
37. The device of claim 35, wherein the device is a L-VAD and is
sized and shaped for positioning between the aortic valve and the
aortic arch of the patient using the device.
38. The device of claim 35, wherein the device is a R-VAD and is
sized and shaped for positioning between the pulmonary valve and
the bifurcation of the pulmonary trunk of the patient using the
device.
39. A system for assisting cardiac ventricular function, the system
comprising a hydraulic pumping assembly and a cardiac ventricular
assist device in fluid communication with the hydraulic pumping
assembly, wherein the ventricular assist device comprises: a. a
first magnet having an open center and formed of high
ferromagnetic-constant material; b. a first vessel surrounding the
first magnet and defining a space in fluid communication with the
blood flow output great vessel of the diseased ventricle of the
patient using the device, the first magnet being movable within the
first vessel in substantially fluid-tight relation thereto; c. a
second magnet formed of high ferromagnetic-constant material and
being in magnetic communication with the first magnet, so that the
respective magnetic fluxes of the first magnet and the second
magnet affect each other, so that the first magnet and the second
magnet are biased toward and tend to lock to one another, to
thereby move in the same direction as one another; d. a second
vessel encasing the second magnet and defining a space, the second
magnet being movable within the space in substantially fluid-tight
relation to the second vessel, the space defined by the second
vessel being in fluid communication with a hydraulic pump for
actuating the second magnet; and e. an one-way valve connected to
the first magnet, the one-way valve being movable with the first
magnet, and adapted to cause movement of blood from the diseased
ventricle to and into the great vessel associated with that
diseased ventricle.
40. The system of claim 39, wherein the ventricular assist device
is a L-VAD and is sized and shaped for positioning between the
aortic valve and the aortic arch of the patient using the
device.
41. The device of claim 39, wherein the ventricular assist device
is a R-VAD and is sized and shaped for positioning between the
pulmonary valve and the bifurcation of the pulmonary trunk of the
patient using the device.
42. A system for assisting cardiac ventricular function, the system
comprising a hydraulic pumping assembly and a cardiac ventricular
assist device in fluid communication with the cardiac ventricular
assist device, wherein the ventricular assist device comprises: a.
a first annular magnet (54) formed of high ferromagnetic-constant
material; b. a first sleeve (68) surrounding the first annular
magnet (54) and defining a space in fluid communication with the
blood flow output great vessel of the diseased ventricle of the
patient using the device, the first annular magnet (54) being
longitudinally and reciprocally slideable within the first sleeve
in substantially fluid-tight relation thereto; c. a second annular
magnet (44) formed of high ferromagnetic-constant material and
sized and shaped for placement exterior of the first sleeve (68),
the second annular magnet (44) being disposed coaxially in relation
to and in magnetic communication with the first annular magnet
(54), so that the respective magnetic fluxes of the first magnet
and the second magnet affect each other, so that the first annular
magnet and the second annular magnet are biased toward and tend to
lock to one another, and to thereby move in the same direction as
one another; d. a second sleeve (72) encasing the second annular
magnet (44), the second annular magnet being longitudinally and
reciprocally slideable between the first sleeve (68) and the second
sleeve in substantially fluid-tight relation to the first sleeve
and the second sleeve, and the second sleeve (72) defining an
annular space (86) radially outwardly of the first sleeve (68) for
longitudinal travel therein of the second annular magnet (44), the
annular space (86) being in fluid communication with a hydraulic
pump for actuating the second annular magnet (44); and e. an
one-way valve (70) connected to the first annular magnet (54) and
disposed transversely in relation to the longitudinal axis of the
first annular magnet (54), the one-way valve being movable with the
first magnet, and closed when moving away from the origin of the
valve annulus when the device is in normal use position, to thereby
cause blood of the patient to move out of a diseased ventricle and
toward the great vessels associated with that diseased ventricle as
the first annular magnet (54) moves in a direction toward the great
vessels of the diseased ventricle due to magnetic flux of the
second annular magnet, the one-way valve further being adapted to
be open when moving in a direction away from the great vessels of
the diseased ventricle, to thereby permit blood of the patient to
flow through the one-way valve into the space defined by the first
sleeve (68) when the second annular magnet moves away from the
great vessels of the diseased ventricle.
43. The system of claim 42, wherein the entire device (58, 74) is
of sufficiently small size and weight for placement in normal
operative position between the valve annulus and the great vessels
of the diseased ventricle of a patient using the device, to assist
or replace the function of at least a portion of the diseased
native heart.
44. A system for assisting cardiac ventricular function, the system
comprising a hydraulic pumping assembly and a cardiac ventricular
assist device (VAD) in fluid communication with the hydraulic
pumping assembly, wherein the hydraulic pumping assembly comprises:
a. an encapsulated hydraulic pump having: a pumping chamber for
retaining hydraulic fluid therein, the pumping chamber having
opposed first and second ends; at least one electromagnetic coil
surrounding the pumping chamber; a substantially solid high
ferromagnetic-constant magnet disposed longitudinally, slideably
and reciprocally within the pumping chamber to act as a piston for
driving hydraulic fluid within the pumping chamber in response to
signals from a battery/controller assembly; b. a fluid line having
a first end and a second end, the first end of the fluid line being
connected to and in fluid communication with an the first end of
the pumping chamber and the second end of the fluid line being
connected to and in fluid communication with the second end of the
pumping chamber, the VAD being in fluid communication with the
fluid line at a point on the fluid line after the point of
connection of the check valve and before the connection of the
second end of the fluid line and the second end pump chamber; and
c. a battery/controller assembly operatively connected to the check
valve and to the at least one electromagnetic coil to provide
electric power and control signals to the pump, the battery
controller assembly in electrical communication with the native
heart of the patient using the system, to thereby receive signals
corresponding to physiological parameters from the native
heart.
45. The system of claim 44, wherein the hydraulic pumping assembly
further comprises a first end cap and a second end cap connected at
opposed first and second ends of the pumping chamber, the first end
cap and the second end cap each having an aperture in fluid
communication with a hydraulic fluid line.
46. The system of claim 45, and further comprising a check valve
operatively connected in the fluid line between the connection of
the first end of the fluid line and the second end of the fluid
line to the first and second ends of the pumping chamber,
respectively and the fluid line being in fluid communication with
the VAD, after the point of connection of the check valve and
before the connection of the second end of the fluid valve and the
second end cap of the pump cylinder.
47. The system of claim 45, wherein the information received from
the native heart by the battery/controller assembly is at least a
portion of an ECG signal from the patient.
48. The system of claim 45, wherein the information received from
the,native heart by the battery/controller assembly is blood
pressure information.
49. The system of claim 45, wherein the information received from
the native heart by the battery/controller assembly is blood volume
information.
50. The system of claim 45, wherein the at least one electromagnet
coil is three electromagnetic coils disposed longitudinally and
coaxially adjacent to one another along the length of the hydraulic
pump.
51. A system for assisting cardiac ventricular function, the system
comprising a hydraulic pumping assembly and a cardiac ventricular
assist device (VAD) in fluid communication with the hydraulic
pumping assembly, wherein the hydraulic pumping assembly comprises:
a. a hydraulic pump (42) having: at least one electromagnetic coil
(46, 48, 50) encapsulated so as to be fluid-tight, and defining a
pumping chamber for retaining hydraulic fluid (52) therein, the
pumping chamber having first and second opposed ends; a first end
cap (56) and a second end cap (57) connected at opposed first and
second ends of the pumping chamber, respectively, the first end cap
and the second end cap each having an aperture in fluid
communication with a hydraulic fluid line, a substantially solid
high ferromagnetic-constant magnet (40) disposed longitudinally,
slideably and reciprocally within the pumping chamber to act as a
piston for driving hydraulic fluid within the pumping chamber in
response to signals from a batter/controller assembly; b. a fluid
line (59, 60)having a first end and a second end, the first end of
the fluid line being connected to and in fluid communication with
an aperture in the first end cap and the second end of the fluid
line being connected to and in fluid communication with an aperture
in the second end cap, and the VAD being in fluid communication
with the hydraulic pumping assembly at a point on the fluid line
between the first end and the second end of the fluid line; c. a
check valve (84) operatively connected in the fluid line between
the connection of the first end of the fluid line and the second
end of the fluid line to the first and second end caps
respectively, and the fluid line being in fluid communication with
the VAD, after the point of connection of the check valve and
before the connection of the second end of the fluid valve and the
second end cap of the pump cylinder; d. a battery/controller
assembly (65) operatively connected to the check valve and to the
at least one electromagnetic coil to provide electric power and
control signals to the pump, the battery controller assembly in
electrical communication with the native heart of the patient using
the system, to thereby receive electronic information, including at
least portions of ECG signals, blood pressure signals and/or blood
volume signals, from the native heart.
52. The system of claim 51, wherein the battery controller assembly
and the hydraulic pump are of sufficiently small size and weight to
be entirely contained within the abdominal cavity of the patient
using the system and the VAD is of sufficiently small size and
weight to be entirely contained within the chest cavity of the
patient using the system, and the complete system, including all
wires and hydraulic fluid lines, is entirely contained within the
body of the patient using the system, so that there is no part of
the system extending exterior of the skin of a patient using the
system when the system is in normal use position in the
patient.
53. A system for assisting cardiac ventricular function, the system
comprising: a hydraulic pumping assembly and a cardiac ventricular
assist device in fluid communication with the cardiac ventricular
assist device, wherein the ventricular assist device comprises: a.
a first annular magnet (54) formed of high ferromagnetic-constant
material; b. a first sleeve (68) surrounding the first annular
magnet (54) and defining a space in fluid communication with the
blood flow output great vessel of a diseased ventricle of the
patient using the device, the first annular magnet (54) being
longitudinally and reciprocally slideable within the first sleeve
in substantially fluid-tight relation thereto; c. a second annular
magnet (44) formed of high ferromagnetic-constant material and
sized and shaped for placement exterior of the first sleeve (68),
the second annular magnet (44) being disposed coaxially in relation
to and in magnetic communication with the first annular magnet
(54), so that the respective magnetic fluxes of the fist magnet and
the second magnet affect each other, so that the first annular
magnet and the second annular magnet are biased toward and tend to
lock to one another, and to thereby move in the same direction as
one another; d. a second sleeve (72) encasing the second annular
magnet (44), the second annular magnet being longitudinally and
reciprocally slideable between the first sleeve (68) and the second
sleeve in substantially fluid-tight relation to the first sleeve
and the second sleeve, and the second sleeve (72) defining an
annular space (86) radially outwardly of the first sleeve (68) for
longitudinal travel therein of the second annular magnet (44), the
annular space (86) being in fluid communication with a hydraulic
pump for actuating the second annular magnet (44); and e. an
one-way valve (70) connected to the first annular magnet (54) and
disposed transversely in relation to the longitudinal axis of the
first annular magnet (54), the one-way valve being movable with the
first magnet, and closed when moving away from the origin of the
valve annulus when the device is in normal use position, to thereby
cause blood of the patient to move out of a diseased ventricle and
toward the great vessels associated with that diseased ventricle as
the first annular magnet (54) moves in a direction toward the great
vessels of the diseased ventricle due to magnetic flux of the
second annular magnet, the one-way valve further being adapted to
be open when moving in a direction away from the great vessels of
the diseased ventricle, to thereby permit blood of the patient to
flow through the one-way valve into the space defined by the first
sleeve (68) when the second annular magnet moves away from the
great vessels of the diseased ventricle; and further wherein the
hydraulic pumping assembly comprises: a. a hydraulic pump (42)
having: at least one electromagnetic coil (46, 48, 50) encapsulated
so as to be fluid-tight, and defining a pumping chamber for
retaining hydraulic fluid (52) therein, the pumping chamber having
first and second opposed ends; a first end cap (56) and a second
end cap (57) connected at opposed first and second ends of the
pumping chamber, respectively, the first end cap and the second end
cap each having an aperture in fluid communication with a hydraulic
fluid line, a substantially solid high ferromagnetic-constant
magnet (40) disposed longitudinally, slideably and reciprocally
within the pumping chamber to act as a piston for driving hydraulic
fluid within the pumping chamber in response to signals from a
batter/controller assembly; b. a fluid line (59, 60) having.a first
end and a second end, the first end of the fluid line being
connected to and in fluid communication with an aperture in the
first end cap and the second end of the fluid line being connected
to and in fluid communication with an aperture in the second end
cap, and the VAD being in fluid communication with the hydraulic
pumping assembly at a point on the fluid line between the first end
and the second end of the fluid line; c. a check valve (84)
operatively connected in the fluid line between the connection of
the first end of the fluid line and the second end of the fluid
line to the first and second end caps respectively, and the fluid
line being in fluid communication with the VAD, after the point of
connection of the check valve and before the connection of the
second end of the fluid valve and the second end cap of the pump
cylinder; and d. a battery/controller assembly (65) operatively
connected to the check valve and to the at least one
electromagnetic coil to provide electric power and control signals
to the pump, the battery controller assembly in electrical
communication with the native heart of the patient using the
system, to thereby receive electronic information, including at
least portions of ECG signals, blood pressure signals and/or blood
volume signals, from the native heart.
54. The system of claim 53, wherein the entire device (58, 74) is
of sufficiently small size and weight for placement in normal
operative position between the valve annulus and the great vessels
of the diseased ventricle of a patient using the device, to assist
or replace the function of at least a portion of the diseased
native heart.
55. The system of claim 53, wherein the battery controller assembly
and the hydraulic pump are of sufficiently small size and weight to
be entirely contained within the abdominal cavity of the patient
using the system and the VAD is of sufficiently small size and
weight to be entirely contained within the chest cavity of the
patient using the system, and the complete system, including all
wires and hydraulic fluid lines, is entirely contained within the
body of the patient, so that there is no part of the system
extending exterior of the skin of a patient using the system when
the system is in normal use position in the patient.
56. A BI-VAD assembly to assist the function of both the right and
left cardiac ventricles simultaneously, the BI-VAD assembly
comprising: a. a L-VAD disposed between the aortic valve and the
aortic arch, to thereby permit blood to move from the left
ventricle of the native heart through the aortic valve and into the
L-VAD of a patient using the system, the L-VAD pumping blood into
the aortic arch; and b. a R-VAD disposed between the pulmonary
valve and the bifurcation of the pulmonary trunk in normal use
position in a patient using the BI-VAD, to thereby permit blood to
move from the right ventricle of the patient through the pulmonary
valve and into the R-VAD as the R-VAD pumps blood into the
bifurcation of the pulmonary arteries of the patient.
57. A BI-VAD assembly (77) to assist the function of both the right
and left cardiac ventricles simultaneously, the BI-VAD assembly
comprising: a. a L-VAD and a R-VAD; b. the L-VAD being sized and
shaped for positioning between the aortic valve and the aortic arch
of the patient using the device; the L-VAD comprising: a. a first
magnet having an open center and formed of high
ferromagnetic-constant material; b. a first vessel surrounding the
first magnet and defining a space in fluid communication with the
aortic arch of the patient using the device, the first magnet being
movable within the first vessel in substantially fluid-tight
relation thereto; c. a second magnet formed of high
ferromagnetic-constant material and being in magnetic communication
with the first magnet, so that the respective magnetic fluxes of
the first magnet and the second magnet affect each other, so that
the first magnet and the second magnet are biased toward and tend
to lock to one another, to thereby move in the same direction as
one another; d. a second vessel encasing the second magnet and
defining a space, the second magnet being movable within the space
in substantially fluid-tight relation to the second vessel, the
space defined by the second vessel being in fluid communication
with a hydraulic pump for actuating the second magnet; and e. an
one-way valve connected to the corresponding first magnet of the
L-VAD, the one-way valve being movable with the first magnet of the
L-VAD and closed when moving in a direction toward the aortic arch
when the device is in normal use position in a patient using the
device, to thereby cause blood of the patient to push through the
aortic arch as the first magnet moves toward the aortic arch of the
patient when the second magnet is actuated to move toward the
aortic arch, the one-way valve in the L-VAD further being open when
moving away from the aortic arch, to thereby permit blood of the
patient to flow through the one-way valve of the L-VAD into the
space defined by the first vessel when the second magnet of the
L-VAD is moved away from the aortic arch; and c. the R-VAD being
sized and shaped for positioning between the pulmonary valve and
the bifurcation of the pulmonary trunk of the patient using the
device and connected to the L-VAD; the R-VAD comprising: a. a first
magnet having an open center and formed of high
ferromagnetic-constant material; b. a first vessel surrounding the
first magnet and defining a space in fluid communication with the
bifurcation of the pulmonary arteries of the patient using the
device, the first magnet being movable within the first vessel in
substantially fluid-tight relation thereto; c. a second magnet
formed of high ferromagnetic-constant material and being in
magnetic communication with the first magnet, so that the first
magnet and the second magnet are biased toward and tend to lock to
one another, to thereby move in the same direction as one another;
d. a second vessel encasing the second magnet and defining a space,
the second magnet being movable within the space in substantially
fluid-tight relation to the second vessel, the space defined by the
second vessel being in fluid communication with a hydraulic pump
for actuating the second magnet; and e. an one-way valve being
movable with the first magnet of the R-VAD and closed when moving
toward the bifurcation of the pulmonary arteries when the R-VAD is
in normal use position in a patient using the assembly, to thereby
cause blood of the patient to push through the bifurcation of the
pulmonary arteries as the first magnet of the R-VAD moves toward
such bifurcation when the second magnet of the R-VAD is actuated to
move toward the bifurcation, the one-way valve of the R-VAD further
being open when moving away from the bifurcation of the pulmonary
arteries, to thereby permit blood of the patient to flow through
the one-way valve of the R-VAD into the space defined by the first
vessel when the second magnet of the R-VAD is moved away from the
bifurcation of the pulmonary arteries.
58. A system for assisting cardiac ventricular function
simultaneously in both diseased ventricles of the native heart of
a.patient using the system, the system comprising at least one
hydraulic pumping assembly and two cardiac ventricular assist
devices in fluid communication with the at least one hydraulic
pumping assembly, wherein the ventricular assist devices comprise:
a. a L-VAD disposed between the aortic valve and the aortic arch of
the patient, to thereby permit blood to move from the left
ventricle of the native heart through the aortic valve and into the
L-VAD in a patient using the system, the L-VAD pumping blood into
the aortic arch; and b. a R-VAD disposed between the pulmonary
valve and bifurcation of the pulmonary trunk in normal use position
in a patient using the BI-VAD, to thereby permit blood to move from
the right ventricle of the patient through the pulmonary valve and
into the R-VAD as the R-VAD pumps blood into the bifurcation of the
pulmonary arteries of the patient; and wherein the at least one
hydraulic pumping assembly comprises: a. a hydraulic pump having:
an encapsulated pumping chamber for retaining hydraulic fluid
therein, the pumping chamber having opposed first and second ends;
at least one electromagnetic coil surrounding the pumping chamber;
and a substantially solid high ferromagnetic-constant magnet
disposed longitudinally, slideably and reciprocally within the
pumping chamber to act as a piston for driving hydraulic fluid
within the pumping chamber in response to signals from a
battery/controller assembly; b. a fluid line having a first end and
a second end, the first end of the fluid line being connected to
and in fluid communication with an the first end of the pumping
chamber and the second end of the fluid line being connected to and
in fluid communication with the second end of the pumping chamber,
the L-VAD and the R-VAD being in fluid communication with the fluid
line at a point on the fluid line after the point of connection of
the check valve and before the connection of the second end of the
fluid line and the second end pump chamber; and c. a
battery/controller assembly operatively connected to the check
valve and to the at least one electromagnetic coil to provide
electric power and control signals to the pump, the battery
controller assembly in electrical communication with the native
heart of the patient using the system, to thereby receive signals
corresponding to physiological parameters from the native
heart.
59. The system of claim 58, wherein the at least one hydraulic
pumping assembly is two hydraulic pumping assemblies and the L-VAD
and the R-VAD are each in fluid communication with a separate one
of the two hydraulic pumping assemblies.
60. A system for completely replacing cardiac ventricular function
in a diseased native heart, the system comprising: a. a hydraulic
pumping system; and b. a BI-VAD assembly having a L-VAD and a
R-VAD, the L-VAD and the R-VAD having sufficient stroke volumes to
supply the total cardiac blood flow output for the diseased native
heart of a patient using the system, the L-VAD being disposed to at
least partly replace the diseased left ventricle of the native
heart of the patient and the R-VAD being disposed in normal use
position to at least partly replace the diseased right ventricle of
the native heart of the patient, with the inlet of the R-VAD being
grafted to an artificial heart valve and the outlet of the R-VAD
being grafted into the pulmonary trunk of the patient; wherein the
L-VAD comprises: a. a first magnet having an open center and formed
of high ferromagnetic-constant material; b. a first vessel
surrounding the first magnet and defining a space in fluid
communication with the aortic arch of the patient using the device,
the first magnet being movable within the first vessel in
substantially fluid-tight relation thereto; c. a second magnet
formed of high ferromagnetic-constant material and being in
magnetic communication with the first magnet, so that the first
magnet and the second magnet are biased toward and tend to lock to
one another, to thereby move in the same direction as one another;
d. a second vessel encasing the second magnet and defining a space,
the second magnet being movable within the space in substantially
fluid-tight relation to the second vessel, the space defined by the
second vessel being in fluid communication with a hydraulic pump
for actuating the second magnet; and e. an one-way valve connected
to the corresponding first magnet of the L-VAD , the one-way valve
being movable with the first magnet of the L-VAD and closed when
moving in a direction toward the aortic arch when the device is in
normal use position in a patient using the device, to thereby cause
blood of the patient to push through the aortic arch as the first
magnet moves toward the aortic arch of the patient when the second
magnet is actuated to move toward the aortic arch, the one-way
valve in the L-VAD further being open when moving away from the
aortic arch, to-thereby permit blood of the patient to flow through
the one-way valve of the L-VAD into the space defined by the first
vessel when the second magnet of the L-VAD is moved away from the
aortic arch; and wherein the R-VAD comprises: a. a first magnet
having an open center and formed of high ferromagnetic-constant
material; b. a first vessel surrounding the first magnet and
defining a space in fluid communication with the bifurcation of the
pulmonary arteries of the patient using the device, the first
magnet being movable within the first vessel in substantially
fluid-tight relation thereto; c. a second magnet formed of high
ferromagnetic-constant material and being in magnetic communication
with the first magnet, so that the respective magnetic fluxes of
the first magnet and the second magnet affect each other, so that
the first magnet and the second magnet are biased toward and tend
to lock to one another, to thereby move in the same direction as
one another; d. a second vessel encasing the second magnet and
defining a space, the second magnet being movable within the space
in substantially fluid-tight relation to the second vessel, the
space defined by the second vessel being in fluid communication
with a hydraulic pump for actuating the second magnet; and e. an
one-way valve being movable with the first magnet of the R-VAD and
closed when moving toward the bifurcation of the pulmonary arteries
when the R-VAD is in normal use position in a patient using the
assembly, to thereby cause blood of the patient to push through the
bifurcation of the pulmonary arteries as the first magnet of the
R-VAD moves toward such bifuircation when the second magnet of the
R-VAD is actuated to move toward the bifurcation, the one-way valve
of the R-VAD further being open when moving away from the
bifurcation of the pulmonary arteries, to thereby permit blood of
the patient to flow through the one-way valve of the R-VAD into the
space defined by the first vessel when the second magnet of the
R-VAD is moved away from the bifurcation of the pulmonary arteries;
wherein the at least one hydraulic pumping assembly comprises: a. a
hydraulic pump having: an encapsulated pumping chamber for
retaining hydraulic fluid therein, the pumping chamber having
opposed first and second ends; at least one electromagnetic coil
surrounding the pumping chamber; and a substantially solid high
ferromagnetic-constant magnet disposed longitudinally, slideably
and reciprocally within the pumping chamber to act as a piston for
driving hydraulic fluid within the pumping chamber in response to
signals from a battery/controller assembly; b. a fluid line having
a first end and a second end, the first end of the fluid line being
connected to and in fluid communication with an the first end of
the pumping chamber and the second end of the fluid line being
connected to and in fluid communication with the second end of the
pumping chamber, the L-VAD and the R-VAD being in fluid
communication with the fluid line at a point on the fluid line
after the point of connection of the check valve and before the
connection of the second end of the fluid line and the second end
pump chamber; and c. a battery/controller assembly operatively
connected to the check valve and to the at least one
electromagnetic coil to provide electric power and control signals
to the pump, the battery controller assembly in electrical
communication with the native heart of the patient using the
system, to thereby receive signals corresponding to physiological
parameters from the native heart.
61. A system for assisting cardiac ventricular function, the system
comprising: a. a ventricular assist device (VAD) having: an
open-centered magnet, at least one encapsulated electromagnetic
coil in magnetic communication with the open-centered magnet; to
thereby drive the magnet; and a one-way valve connected to the
open-centered magnet, the one-way valve being movable with the
magnet, and adapted to cause movement of blood from the diseased
ventricle to and into the great vessel associated with the diseased
ventricle; b. a battery/controller assembly operatively connected
to the at least one electro-magnetic coil for energizing same and
connected to the sino-atrial node of the patient when the system is
in normal operative position in the patient to thereby provide
signals to the VAD from the sino-atrial node to activate the at
least one electromagnetic coil to optimally complement the function
of the diseased ventricle of the patient's native heart.
Description
BACKGROUND OF THE INVENTION
[0001] 1. Field of the Invention
[0002] This invention relates generally to the field of mechanical
cardiac pumping devices, and, more particularly, to a ventricular
assist device (VAD) and a total artificial heart (TAH) device and
method of using same. More specifically, this invention relates to
a VAD and a TAH that are optimized by the new method to produce
customized pulsatile blood flow mimicking that of the healthy
native heart for each individual patient case.
[0003] 2. Description of Related Art
Introduction
[0004] Some medical studies indicate: a) 400,000 new cases of
congestive heart failure are diagnosed annually in the United
States; b) a mortality rate of 75 percent in men and 62 percent in
women; c) standard medical therapies benefit only a limited
percentage of patients with ventricular dysfunction; and d) from
17,000 to 66,000 patients per year, in the United States alone, may
benefit from a permanent implantable blood pump. Presently,
potential cardiac transplant recipients with hemodynamic compromise
(inadequate perfusion of the systemic circulation by the native
heart) sometimes receive temporary mechanical circulatory support
as a "bridge" to permit them to survive until cardiac
transplantation is possible. It is foreseen that some day
mechanical blood pumps will provide a cost-effective alternative to
either cardiac transplantation or long term medical management of
patients. It is to this end that the devices and methods described
herein have been developed.
[0005] It is to be understood that for purposes of this document a
"ventricular-assist device (VAD)" is a mechanical blood pump that
assists a diseased native heart to circulate blood in the body, and
a "total artificial heart (TAH)" is another type of mechanical
blood pump that replaces the native heart and provides all of the
blood pumping action in the body.
[0006] In order for a VAD to function optimally, it must both
complement the diseased native heart and make the combined output
of the VAD and native diseased heart emulate the pumping action of
the natural healthy human heart. That is, it should provide
pulsatile flow similar to that of the healthy heart. In order for a
TAH to function optimally, it must mimic the pulsatile pumping
action of the natural healthy human heart. In either case, the
device must be sized such that it fits within the required areas in
the patient's body. In order to minimize the size of the power
supply portion of the device, each device (VAD or TAH) must use as
little energy and as little power as possible to accomplish the
required function. Thus, there is a need for bio-emulating
efficient pump (BEEP) systems for VAD and TAH applications.
[0007] It is known that VADs can be implanted to assist a
functioning heart that does not have adequate pumping capability.
Often, however, residual cardiac function is not taken into account
in the design of such devices, resulting in less than optimal
effects. What is needed is a bio-emulating efficient pump (BEEP)
system, which works in concert with the native human heart. The new
VAD device and system and optimization procedure described herein
utilize patient specific information concerning residual cardiac
output to optimize the pumping action provided for each individual
patient, thereby providing such a BEEP system. The TAH device and
optimization procedure described in this document optimize the
pumping function provided for each individual patient, thereby
providing such a BEEP system which is customized for each such
patient.
Known Heart Pump Devices
[0008] Previously, a number of devices were developed for blood
pumping. Highly specialized pumps have been used to completely
replace a biological heart which has been surgically removed. Such
known heart pumps may be temporary, or permanently implantable.
Temporary heart pump devices usually involve either: 1) an attempt
to augment a compromised native heart while it recovers from
surgery or some other short-term problem; or 2) use of the device
as a "bridge" to extend the life of a patient by temporarily
replacing the native heart until a suitable donor heart can be
found for cardiac transplantation.
[0009] Many types of permanently implantable heart pumps have been
proposed and several have been developed. Because the left
ventricle of the heart, which pumps blood to the entire body except
for the lungs, becomes diseased far more commonly than the right
ventricle (which pumps blood only to the lungs), most heart pumps
have been developed to assist or replace the left ventricle. Fewer
pumps have been proposed, tested, and used for bi-ventricular
function (i.e. assisting both the left and right ventricles).
[0010] Known mechanical blood pumps can be roughly divided into
three major categories: a. pulsatile sacks; b. reciprocating
piston-type pumps; and c. pumps with axial or centrifugal
impellers. Each category has distinct advantages and
disadvantages.
a. Pulsatile Sacks
[0011] Pulsatile sack devices are the most widely tested and used
implantable blood pumps. These devices employ flexible sacks or
diaphragms which are compressed and released in a periodic manner
to cause pulsatile flow of blood. Sack or diaphragm pumps are
subject to fatigue failure of compliant elements. They are
generally used as temporary heart-assist devices, and they are
mechanically and functionally different from the present invention
described hereafter.
[0012] The intra-aortic balloon (IAB) counter-pulsation device, a
pulsatile sack device, is readily available. It is a
catheter-mounted intra-vascular device designed to improve the
balance between myocardial oxygen supply and demand. The first
successful clinical application of the balloon was reported by
Kantrowitz et al. in 1968. The IAB is positioned in the thoracic
aorta and set to inflate at the dicrotic notch of the atrial
pressure waveform when monitoring aortic pressure. The diastolic
rise in aortic pressure augments coronary blood flow and myocardial
oxygen supply. The IAB is deflated during the isovolumetric phase
of left ventricular contraction. The reduction in the afterload
component of cardiac work decreases peak left ventricular pressure
and myocardial oxygen consumption. These units are not portable and
are limited to in-hospital critical care use only. Use of the IAB
is now a standard form of therapy for a variety of patients with
cardiovascular disease, primarily reserved for patients with
deteriorating heart function while awaiting revascularization
procedure. In 1993, nearly 100,000 IABs were inserted in the United
States alone.
[0013] Another example of a pulsatile sack device is the
Abiomed.TM. BVS.RTM. device (Abiomed, Inc., Boston, Mass.). It is
an externally placed dual-chamber device that is capable of
providing short term univentricular or biventricular support. It
has pneumatically driven polyurethane blood sacks and it is not
intended for long-term support. Also, U.S. Pat. No.4,888,011 to
Kung and Singh discloses a hydraulically driven dual-sack system;
and U.S. Pat. No. 5,743,845 to Runge discloses a sack-operated
bi-ventricular assist device that balances the flow in the left and
right side of the circulatory system.
b. Reciprocating Piston-Type Pumps
[0014] Several types of implantable blood pumps containing a
piston-like member have been proposed to provide a mechanical
device for augmenting or totally replacing the blood pumping action
of a damaged or diseased heart. For example, the HeartMate.RTM.
(Thermo Cardiosystems, Inc., Woburn, Mass.) is a pneumatically
powered device that is implanted in the left upper quadrant of the
abdomen. A pneumatic air hose exits from the lower half of the
abdominal wall and is attached to a pneumatic power unit. Blood
from the cannulated left ventricular apex empties into a pump, at
which point an external control system triggers pumping. The blood
chamber is pressurized by a pusher plate forcing a flexible plastic
diaphragm upward. This motion propels the blood through an outflow
conduit grafted into the aorta, the main artery supplying the body
with blood. This device is unique in that the textured,
blood-containing surface promotes the formation of a stable
neointima, hence full anticoagulation is not necessary, only
anti-platelet agents are required. This device is designed for left
ventricular support only. It uses trileaflet polyurethane valves.
There is an electrically powered version with percutaneous electric
leads connecting the pump to external batteries.
[0015] The Thoratec.RTM. VAD (Thoratec Laboratories, Pleasanton,
Calif.) is a pneumatically powered device that is placed externally
on the anterior abdominal wall. Cannulas pass through the chest
wall in a manner similar to that of a conventional chest tube. The
device takes blood from the left ventricular apex and returns it to
the aorta. Full systemic anticoagulation is required with this
device. It can be used to support either ventricle and uses tilting
disc type mechanical valves.
[0016] Novacor.RTM. (Cedex, France) produces an electrically driven
device that is implanted in the left upper quadrant of the abdomen
and the electric line and vent tube are passed through the lower
anterior abdominal wall. This system also incorporates a
polyurethane blood sac that is compressed by dual symmetrically
opposed pusher plates. Blood is taken from the left ventricular
apex and returned to the aorta. Full anticoagulation is
required.
[0017] U.S. Pat. No. 3,842,440 to Karlson discloses an implantable
linear motor prosthetic heart and control system containing a pump
with a piston-like member which reciprocates in a magnetic field.
The piston includes a compressible chamber in the prosthetic heart
which communicates with the vein or aorta.
[0018] U.S. Pat. Nos. 3,911,897 and 3,911,898 to Leachman, Jr.
disclose heart assist devices controlled in the normal mode of
operation to copulsate and counterpulsate with the heart,
respectively, and produce a blood flow waveform corresponding to
the blood flow waveform of the assisted heart. The heart assist
device is a pump connected serially between the discharge of a
heart ventricle and the vascular system. This pump has cylindrical
inlet and discharge pumping chambers of the same diameter and a
reciprocating piston in one chamber fixedly connected with a
reciprocating piston of the other chamber.
[0019] U.S. Pat. No. 4,102,610 to Taboada et al. discloses a
magnetically operated constant volume reciprocating pump which can
be used as a surgically implantable heart pump or assist. The
reciprocating member is a piston carrying a check valve positioned
in a cylinder.
[0020] U.S. Pat. Nos. 4,210,409 and 4,375,941 to Child disclose a
pump used to assist the pumping action of the heart with a piston
movable in a cylindrical casing in response to magnetic forces. A
tilting-disk type check valve carried by the piston provides for
flow of fluid into the cylindrical casing and restricts reverse
flow.
[0021] U.S. Pat. No. 4,965,864 to Roth discloses a linear motor
using multiple coils and a reciprocating element containing
permanent magnets, driven by microprocessor-controlled power
semiconductors. A plurality of permanent magnets is mounted on the
reciprocating member. U.S. Pat. No. 4,541,787 to DeLong describes a
pump configuration wherein a piston containing a permanent magnet
is driven in a reciprocating fashion along the length of a cylinder
by energizing a sequence of coils positioned around the outside of
the cylinder.
[0022] U.S. Pat. No. 4,610,658 to Buchwald et al. discloses an
implantable fluid displacement peritoneovenous shunt system. The
device is a magnetically driven pump, which can be a reciprocating
diaphragm, or piston type, or rotary pump.
[0023] U.S. Pat. No. 5,089,017 to Young et al. discloses a drive
system for artificial hearts and left ventricular assist devices
comprising one or more implantable pumps driven by external
electromagnets. The pump utilizes working fluid, such as sulfur
hexafluoride to apply pneumatic pressure to increase blood pressure
and flow rate.
[0024] Larson et al. in a series of patents (1997-1999, U.S. Pat.
Nos. 5,879,375; 5,843,129; 5,758,666; 5,722,930; 5,722,429;
5,702,430; 5,693,091; 5,676,651; 5,676,162) describe a piston-type
pump for ventricular assist or total replacement, and associated
driving equipment and power supply. The piston is an artificial
heart valve, with valves that have at least two leaflets, acting as
a check valve and reciprocating in a cylinder. The walls of the
cylinder are a few millimeters thick because they contain the coils
of a linear electric motor that must provide pumping power to the
VAD. Around the artificial heart valve and inside the cylinder is a
hollow cylindrical rare-earth permanent magnet, which is driven by
the linear electric motor. In one embodiment one device is
implanted in series to the aorta (left VAD), or another device is
implanted in series to the pulmonary artery (right VAD), or two
devices are used on both aorta and pulmonary artery (BI-VAD). In a
second embodiment one device replaces the left ventricle, or
another device replaces the right ventricle, or two devices replace
the whole heart.
[0025] Measurements on experimental devices made with hollow pump
cores indicate that such devices are too large to fit in the
available space in the chest cavity in the aorta or pulmonary
artery, due to the size of the coils necessary to drive the device.
For a given volume of blood pumped per stroke, if the length of the
cylinder is restricted such that the device fits lengthwise in the
human body, then the diameter must be increased until the. desired
volume is reached. The outer diameter of the device is severely
restricted by the surrounding tissue, and this leaves little room
available in the diameter for the linear magnet motor. In a
bi-ventricular application, if the axes of the two cylinders are
located in parallel, then even more space is needed due to the
diameters required; and if they are not parallel the magnetic
fields of the two motors introduce additional electromagnetic
losses because the linear magnet motors are not parallel. Even if
the volumetric displacement of the device is reduced in order to
fit in the available space at the expense of throughput, much of
the outside diameter of the device must still be devoted to the
linear motor. However, the most important disadvantage is that the
linear motor is driving an annular magnet containing a one-way
valve, so that the ferromagnetic material can not be in the core
(center) of the motor coils, leading to lower efficiency.
[0026] At the geometric center (axis) of the motor described by
Larson et al. is the artificial valve acting as the piston, and the
blood itself. This structure introduces electromagnetic losses in
the device that make it less desirable than devices that have
ferromagnetic material in the geometric center (axis) of the motor
coils. In addition, voltage propagates at constant velocity from
coil to coil in the linear magnet motor of the Larson et al.
device, and motion of the magnet carrying the artificial heart
valve is coupled to this application of voltage, so that the
application of current in the Larson et al. device is not optimized
to minimize the power required to effect the blood-pumping
action.
c. Pumps with Axial or Centrifugal Impellers
[0027] After pulsatile devices, rotary pumps, having either
centrifugal or axial impellers, are the most widely used and tested
devices. In centrifugal pumps, the blood flow enters axially into a
centrifugal impeller, centrifugal acceleration increases the blood
flow velocity, the flow exits radially, and the flow is
subsequently decelerated to increase blood static pressure in the
diffusion process. Most such centrifugal pumps provide continuous
(non-pulsatile) flow; or flow with a small fluctuating pressure
trace superimposed on a larger steady-pressure component, such as
U.S. Pat. No. 5,928,131 to Prem and U.S. Pat. No. 6,179,773 to Prem
and Kolenik.
[0028] Axial pumps direct blood flow along a cylindrical axis,
which is in a straight (or nearly straight) line with the direction
of the inflow and outflow. The impeller looks like an axial fan, or
propeller, inside a nozzle. The impeller imparts acceleration to
the fluid, and the subsequent deceleration (diffusion) process
increases the blood pressure. Most such axial pumps provide
continuous (non-pulsatile) flow.
[0029] Some types of axial rotary pumps use impeller blades mounted
on a center axle, which is mounted inside a tubular conduit. As the
blade assembly spins, it functions like a fan or an outboard motor
propeller. Another type of axial blood pump, called the "haemopump"
uses a screw-type impeller with a classic screw (also called an
Archimedes screw; also called a helifoil, due to its helical shape
and thin cross-section). In screw-type axial pumps, the screw spins
at very high speed (up to about 10,000 rpm). The entire haemopump
unit is usually less than one centimeter (approximately 0.4 inches)
in diameter. The pump can be passed through a peripheral artery
into the aorta, through the aortic valve, and into the left
ventricle. An external motor and drive unit powers it.
[0030] Axial and centrifugal pumps provide mostly steady
(continuous) flow with an imperceptible high-frequency
low-amplitude pulsatile component. Various mechanisms have been
proposed to convert this practically steady-flow output into
pulsatile flow. However, both axial and centrifugal impeller pumps
introduce rapid acceleration and deceleration forces and large
shear stresses in the blood. As is well known to those with
ordinary skill in the art (Balje, 1981), both types of
turbomachines (axial and centrifugal) are a balanced compromise
between diameter and speed to provide the specified flow rate and
pressure increase. Imposing limits in diameter in order to reduce
shear stresses means that the optimum machine requires a
higher-speed axial component. Imposing speed limits in order to
reduce shear stresses means that the optimum machine requires a
higher-diameter centrifugal component. It is well know to those
with ordinary skill in the art (Wilson and Korakianitis, 1998) that
small impellers that can fit inside the spaces available in the
human body will result in high blood shear, due to the high
operational speed required.
[0031] The Jarvik 2000.RTM. (registered trademark of R. Jarvik, New
York, N.Y.) System consists of a small axial flow pump (about the
size of a C-cell battery) that is placed in the left ventricular
apex and pumps blood into the aorta. It is still currently being
developed and will use external batteries and control electronics
utilizing induction coils to carry the control signals through the
skin. Power is also delivered transcutaneously.
Medical Complications
[0032] According to several medical studies, the above devices are
subject to a number of complications. Insertion of a cannula to
feed a pump can cause damage to the left ventricle. At least 50
percent of patients who are supported for prolonged periods develop
infections, including those associated with pneumatic lines or
electrical leads. Septic emboli may occur, and the mortality rate
is up to 50 percent. VADs may also activate the coagulation
cascade, resulting in thrombi formation. This occurs in the
approximate range of nine to forty-four percent of patients. Stasis
of blood within the pump may lead to thrombus deposition. Right
ventricular failure may occur peri-operatively with placement of a
left VAD. The right heart failure rate may be as high as 33
percent, with one-fifth of those patients dying from the
complication. Rapid recognition of this complication and
implantation of a right VAD may reduce the mortality rate resulting
from right heart failure. Hemorrhage occurs in about 27 to 87
percent of patients who require mechanical ventricular assistance.
Hemorrhage is also related to inflow and outflow cannulae and to
anticoagulation required with the devices.
[0033] One of the most important problems in axial and centrifugal
rotary pumps involves the interface between the edges of the blades
and the blood flow. The outer edges of the blades move at high
speeds and generate high levels of shear. Red blood cells are
particularly susceptible to shear stress damage, as their cell
membranes do not include a reinforcing cytoskeleton to maintain
cell shape. Lysis of red blood cells can result in the release of
cell contents and trigger subsequent platelet aggregation. Lysis of
white blood cells and platelets also occurs upon application of
high shear stress. Even sublytic shear stress leads to cellular
alterations and direct activation and aggregation of platelets.
Rotary pumps generally are not well tolerated by patients for
prolonged periods. In medical tests, animals placed on these units
for a substantial length of time often suffer from strokes, renal
failure, and other organ dysfunction. The device and method of
optimization disclosed herein minimizes the above, and other, known
complications resulting from implantation of either a VAD or a
TAH.
Desirable Pump Characteristics
[0034] In many patients with end stage heart disease, there is
enough residual function left in the native heart to sustain life
in a sedentary fashion, but insufficient reserve for even minimal
activity, such as walking a short distance. This residual function
of the diseased native heart is typically not considered in the
design of most VADs. Most known VADs are designed to assume
complete circulatory responsibility and to receive blood from the
cannulated ventricular apex of the particular ventricle they are
"assisting," in what is commonly called "fill to empty" mode. It
generally takes one or more contractions of the diseased native
ventricle to supply enough blood to the VAD. Once a pre-specified
volume of blood is accumulated in the VAD, then the ejection phase
of the VAD is initiated. Thus, most known VADs operate in this
"fill-to-empty" mode that is in random association with native
heart contraction, and can be installed in parallel to the native
ventricle or in series. These constructions are not considered to
"complement" the native heart, as does the present invention.
[0035] At least some residual cardiac function is present in the
majority of patients who would be candidates for mechanical
circulatory assistance. It is preferable for the natural heart to
continue contributing to the cardiac output even after a mechanical
circulatory device is installed. This points away from the use of
total cardiac replacements and suggests the use of assist devices
whenever possible. However, the use of assist devices also poses a
very difficult problem. In patients suffering from severe heart
disease, temporary or intermittent crises often require artificial
pumps to provide bridging support which is sufficient to entirely
replace ventricular pumping capacity for limited periods of time.
Such requirements arise in the hours or days following a heart
attack or cardiac arrest, or during periods of certain life
threatening arrhythmias. Therefore, there is an important need to
provide a pump and method that can meet a wide spectrum of
requirements by providing two different and distinct pumping
functions, assisting the native heart and total substitute pump
support.
SUMMARY OF THE INVENTION
[0036] The present invention provides a cardiac ventricular-assist
device and method of optimizing any design of VAD or TAH wherein
the amount of power required by the device is minimized to the
extent necessary to complement the cardiac output of the native
heart, and no more. In this manner, the weight and size of the
device are kept within suitable reasonable ranges to permit
placement of the VAD/TAH within the body of the subject patient
using the new device.
[0037] The present invention further provides a VAD and method
wherein the principles of unsteady thermodynamics and fluid
mechanics are used to provide a uniquely optimized pulsatile blood
flow which complements the cardiac output of the individual native
human heart. It is to be understood that throughout this document,
when the terms "optimize" and "complement" are used in reference to
the devices and systems of the present invention, it is meant that
at each heart beat and stroke of the VAD (used here to mean either
the L-VAD, R-VAD, BI-VAD or TAH as described below), several
actions are carefully timed such that:
[0038] a) the native heart is allowed to pump as much blood as it
can on its own before the VAD is activated;
[0039] b) as the blood-ejection phase of the native heart nears
completion, the VAD is energized to provide additional pumping
action;
[0040] c) the additional pumping action reduces the back pressure
in that native ventricle so that the native ventricle pumps more
than it would have pumped unaided;
[0041] d) the timing of the action, length of pumping stroke, and
rate of pumping (stroke displacement versus time and resulting
power input versus time) of the VAD are related to the native heart
ejected blood volume and rhythm in a manner that minimizes power
input to the VAD while meeting physiological constraints;
[0042] e) the optimization processes in d) take into account the
dynamic interaction between the native heart and the VAD; and
[0043] f) the optimization process and the control scheme are
integrated with the resulting changes in blood ejected per heart
beat and heart rate (beats per minute) by the combined action of
the native heart and the VAD.
[0044] Before turning to the Figures, it is considered useful to
provide some introductory material. The present invention,
described below, is distinct from each of the three categories of
mechanical circulatory support devices previously described, and
consolidates the advantages and avoids the disadvantages of each
category. First, it is carefully noted that several of the devices
described in the known art mention that the power input is
"optimized", but they do not describe how this is accomplished. The
optimization method described herein can be applied to all existing
VAD and TAH devise that have been devised to date, or will be
devised in the future.
[0045] The pump of the present invention has ferromagnetic material
as the solid center of the motor coils, thus providing a more
compact arrangement of the electromagnetic fluxes than pumps with
non-ferromagnetic centers, and simultaneously permitting reduction
of electromagnetic losses in use. Ultimately this permits placement
of a device that can pump sufficient volume per stroke at the
outlet of the native ventricles and allows the power supply to be
smaller than was possible with previous cardiac pumping devices.
The remote hydraulic drive and power supply/controller assembly are
located in the abdomen, thus allowing practically all available
space in the vicinity of the heart for use by the device. Power is
transmitted hydraulically from the abdomen to the blood pump in the
vicinity of the heart. Also, electromagnetic losses are not
introduced by the location of the two pumping devices (artificial
heart valves) in non-parallel configuration in the vicinity of the
aorta and pulmonary artery.
[0046] Details of the dynamics of the pumping action of the human
heart have been incorporated for the first time into the design of
the VADs and TAHs in the present invention. Understanding these
details:
[0047] is essential for optimization of the timing of unsteady-flow
events in the heart-pumping cycle;
[0048] directly impacts the optimum geometric shape of the
artificial devices; and
[0049] identifies prerequisite means to minimize shear stresses on
the blood (reducing blood-cell lysis) and optimizing energy flows
(reducing the power input required to produce the required blood
flow and pressure characteristics).
[0050] The adult heart is located between the lungs and is about
the size of a large grapefruit, weighing 0.2 to 0.5 kg (0.44 to 1.1
pounds), depending on the size of the individual. The
cardiovascular system performs two major tasks: it delivers oxygen
and nutrients to body organs; and removes waste products of
metabolism from tissue cells. Its major components are: the heart
(a two-sided biological pump); and the circulatory system of
elastic blood vessels (veins and arteries) that transport blood. As
an example, the heart of a 70 kg (154 pounds) human circulates
about 6 kg (13.2 pounds), or 6 L (6.34 qt.s), of blood.
[0051] The human heart is divided into four chambers: the right
atrium and right ventricle; and the left atrium and left ventricle.
The walls of the chambers are made of a special muscle called the
myocardium that contracts rhythmically under electric stimulation.
The left and right atria are separated from each other by the
atrial septum; and the left and right ventricles are separated from
each other by the ventricular septum.
[0052] In the circulatory system, blood returns by the venous
system from the body and enters the heart through the right atrium,
then subsequently blood enters the right ventricle. Each time the
right ventricle contracts, it propels this blood (low in oxygen
content) into the lungs, where it is enriched with oxygen.
Pulmonary veins return the-blood to the left atrium, then
subsequently the blood enters the left ventricle. The left
ventricle, which traditionally has been considered as the main
pumping instrument of the heart, ejects the blood through the main
artery, the aorta, to supply oxygenated blood to the various organs
of the body. The organs use the oxygen and with capillary action
between the arterioles and the venules return the blood to the
venous system and the right atrium. The pumping action of the left
and right side of the heart generates pulsatile flow and pressure
on the aorta and pulmonary artery, discussed further below.
[0053] Blood is kept flowing in this pulsatile cycle by a system of
four one-way valves in the heart, each closing an inlet or outlet
in one of the heart's four chambers at the appropriate time in the
cardiac cycle. The valve system helps maintain a pressure
difference between the right and left sides of the heart. The
aortic valve and the pulmonary valve each have three tissue cusps
(leaflet flaps), referred to as "semilunar valves" because of the
crescent shape of these cusps. The tricuspid and mitral valves
separate the atria from the ventricles. The mitral valve has two
cusps and the tricuspid valve has three cusps. In addition, the
cusps have thin chords of fibrous tissue (chordae tendineae), which
tether the valves to the ventricular walls. When the ventricles
contract, small muscles in their walls (papillary muscles) restrict
closure of the mitral and tricuspid valve leaflets, preventing them
from overextending.
[0054] Electric currents control the,pumping motion of the heart.
The currents originate in the sinus node (the heart's natural
pacemaker), a microscopic bundle of specialized cells located in
the superior portion of the atria. The currents travel through a
network of conducting fibers to the atrioventricular or AV node,
the bundle of His, and the Purkinje fibers. The electric currents
cause impulses that are transmitted and propagate in a wave fashion
through the muscle fibers of the left and right atria to the
atrioventricular node, located on the juncture between the right
and left sides of the heart where the right atrium and right
ventricle meet. From the atrioventricular node, they travel along
the bundle of His and the Purkinje fibers through the muscles of
the right and left ventricles. Most currents in the heart are less
than a millionth of an Ampere, but they exert a powerful influence
on the heart muscle.
[0055] The new VAD utilizes electromagnetic coils to drive a
high-ferromagnetic-constant driving magnet in a reciprocating
fashion so as to act as a piston for hydraulic fluid. The resultant
movement of hydraulic fluid through the system in turn moves
another magnet, which is annular, and which also drives in a
reciprocating fashion. The movement of the driven annular magnet in
turn moves still another magnet, an annular valve seat magnet,
which supports a one-way valve. This valve seat magnet is located
inside the annular driven magnet, the two magnets sharing a common
center axis, hence coupling them together. The one-way valve pushes
blood through the ascending aorta of the heart when the valve is
pushed forward, and allows blood to flow freely past when the
one-way valve is moved backward.
[0056] The present invention provides a ventricular-assist device
and method for optimizing same that can be utilized to assist
either the left ventricle (L-VAD) or right ventricle (R-VAD) of the
native human heart or, if necessary, to assist both cardiac
ventricles (BI-VAD). The L-VAD, R-VAD and BI-VAD devices all
utilize principles of unsteady fluid mechanics to provide a
uniquely individualized optimized pulsatile blood flow for each
particular patient.
[0057] In an alternative embodiment, a total artificial heart (TAH)
device that utilizes the principles of unsteady fluid mechanics
provides a uniquely individualized optimized pulsatile blood flow
for each particular patient. The optimized pulsatile blood flow
mimics that of the native heart while simultaneously minimizing the
power required to drive the TAH device.
[0058] Accordingly, it is among the goals of the present invention
to provide a cardiac pump (VAD or TAH) device and system, and
method for controlling and operating same which permit customized,
optimized "assist" or "total" ("complete") cardiac pumping support
for an indefinite period of time. Under appropriate conditions, the
new VAD acts synergistically with the native heart to provide a
seamless augmentation to the otherwise suboptimal output of the
diseased native heart. This allows the new pump device (VAD) to
take advantage of the natural, non-hemolytic pumping action of the
native heart to the fullest extent possible to minimize red blood
cell lysis, and to reduce mechanical stress on the VAD system pump,
requiring less volume, less energy, and hence allowing longer pump
life and longer battery life.
[0059] Accordingly, in furtherance of the above objects and goods,
the present invention is, briefly, a method of optimizing a
mechanical cardiac pumping device includes modeling the physical
system, or portions thereof, of the patient who will receive the
mechanical cardiac pumping device and identifying an operating
condition of the native heart to which the device will respond. The
model is used to determine the required blood volume to be ejected
from the device and an initial estimate of the power required to be
provided to the mechanical cardiac pumping device is provided in
order to provide the required ejected blood volume. The resultant
ejected blood volume is evaluated with data obtained from the model
and the estimate of the power requirement is then updated. The
above steps are iteratively performed until the power required to
obtain the necessary ejected blood volume is identified. Possible
variations of power and pumping rate that allow the mechanical
cardiac pumping device to provide the required volume are
determined and the variation that best matches the physiological
constraints of the patient and minimizes the power required by the
mechanical cardiac pumping device is selected. The steps are
iteratively performed until the mechanical cardiac pumping device
is optimized to respond to each desired operating condition of the
native heart.
[0060] The mechanical system for accomplishing the new method is,
briefly, a system for assisting cardiac ventricular function, the
system including a hydraulic pumping assembly and a cardiac
ventricular assist device (VAD) in fluid communication with the
hydraulic pumping assembly, wherein the hydraulic pumping assembly
includes an encapsulated hydraulic pump having a pumping chamber
for retaining hydraulic fluid therein. The pumping chamber has
opposed first and second ends and at least one electromagnetic coil
surrounding the pumping chamber. A substantially solid high
ferromagnetic-constant magnet is disposed longitudinally, slideably
and reciprocally within the pumping chamber to act as a piston for
driving hydraulic fluid within the pumping chamber in response to
signals from a battery/controller assembly. A fluid line has a
first end and a second end. The first end of the fluid line is
connected to and in fluid communication with the first end of the
pumping chamber and the second end of the fluid line is connected
to and in fluid communication with the second end of the pumping
chamber. The VAD is in fluid communication with the fluid line at a
point on the fluid line after the point of connection of the check
valve and before the connection of the second end of the fluid line
and the second end pump chamber. A battery/controller assembly is
operatively connected to the check valve and to the at least one
electromagnetic coil to provide electric power and control signals
to the pump. The battery controller assembly is in electrical
communication with the native heart of the patient using the
system, to thereby receive signals corresponding to physiological
parameters from the native heart for transfer to the VAD.
[0061] The new VAD device is, briefly, a device to assist the
function of a cardiac ventricle, the device having a first magnet
with an open center and formed of high ferromagnetic-constant
material. A first vessel of the device surrounds the first magnet
and defines a space in fluid communication with the blood flow
output great vessel associated with the diseased ventricle of a
patient using the device, the first magnet being movable within the
first vessel in substantially fluid-tight relation thereto. A
second magnet is formed of high ferromagnetic-constant material in
magnetic communication with the first magnet, so that the magnetic
fluxes of the first magnet and the second magnet affect each other,
so that the first magnet and the second magnet are biased toward
and tend to lock to one another, to thereby move in the same
direction as one another. A second vessel encases the second magnet
and defines a space and is movable within the space in
substantially fluid-tight relation to the second vessel, the space
being defined by the second vessel being in fluid communication
with a hydraulic pump for actuation the second magnet. A one-way
valve is connected to the first magnet, the one-way vale being
movable with the first magnet, and adapted to cause movement of
blood from the diseased ventricle to and into the great vessel
associated with the diseased ventricle.
[0062] These and other advantageous features of the present
invention will be in part apparent and in part pointed out herein
below.
BRIEF DESCRIPTION OF THE DRAWINGS
[0063] FIG. 1 is a schematic view generally identifying a
bio-emulating efficient pump (BEEP) system. The Figure specifically
illustrates the left ventricular-assist device (L-VAD) embodiment
of a BEEP system at the beginning of the blood-pumping stroke.
[0064] FIG. 2 is a schematic view of the L-VAD embodiment of a BEEP
system of FIG. 1, wherein the system is near the middle of the
blood-pumping stroke.
[0065] FIG. 3 is a schematic view of the L-VAD embodiment of a BEEP
system of FIG. 1, wherein the system is at the beginning of the
return stroke.
[0066] FIG. 4 is a schematic view of the L-VAD embodiment of a BEEP
system of FIG. 1, wherein the system is near the middle of the
return stroke.
[0067] FIG. 5 is a cross-sectional view of the hydraulic pump of
the BEEP system of FIG. 1, along line 5-5.
[0068] FIG. 6 is a cross-sectional view of the L-VAD of the BEEP
system of FIG. 1, along line 6-6.
[0069] FIG. 7 is a schematic concept illustration of the human
heart illustrating the location of an L-VAD in place of at least
part of the ascending aorta.
[0070] FIG. 8 a schematic sectional view of a human torso O,
illustrating the location of the main components of an L-VAD
embodiment of BEEP system 35 in the human body. The L-VAD is shown
in place of the ascending aorta, and the hydraulic pump and
battery/controller assembly are illustrated in the abdominal
cavity.
[0071] FIG. 9 is a concept illustration of the human heart
illustrating location of proximity sensors embedded in the
endocardial surface of the left and right ventricles, and mounted
on the pericardium.
[0072] FIG. 10 is a concept illustration of the human heart
illustrating the KG diaphragm in late diastole.
[0073] FIG. 11 is a concept illustration of the human heart
illustrating the KG diaphragm in early systole.
[0074] FIG. 12 is a concept illustration of the human heart
illustrating the KG diaphragm in late systole.
[0075] FIG. 13 is a concept illustration of the human heart
illustrating the KG diaphragm in early diastole.
[0076] FIG. 14 is a graph illustrating typical pressure-volume
diagrams of a native healthy heart and a native diseased heart.
[0077] FIG. 15 is a left-ventricle pressure versus time diagram of
a native healthy heart and a native diseased heart.
[0078] FIG. 16 is a graph illustrating the relationship between the
travel of the piston of the present device and the residual cardiac
output provided by the native diseased heart.
[0079] FIG. 17 is a series of graphs comparing the position and
power requirements of a prior art pumping system and the present
BEEP system with respect to a typical electro cardio gram (ECG)
trace.
[0080] FIG. 18 illustrates the location of three coils of one
embodiment of the BEEP system and the corresponding current flow
sequence in the coils.
[0081] FIG. 19 illustrates the location of three electromagnetic
coils in one embodiment of the BEEP system and the corresponding
current flow sequence in the coils when only two of the coils are
used to move the piston.
[0082] FIG. 20 is a concept illustration of the human heart
illustrating the location of a right ventricular-assist device
(R-VAD) embodiment of the BEEP system.
[0083] FIG. 21 a schematic view of the human torso illustrating the
location of the main components of an R-VAD embodiment of a BEEP
system in the human body.
[0084] FIG. 22 is a concept illustration of the human heart
illustrating the location of a bi-ventricular-assist device
(BI-VAD) embodiment of the BEEP system.
[0085] FIG. 23 a schematic view of the human torso illustrating the
location of the main components of a BI-VAD embodiment of a BEEP
system in the human body
[0086] FIG. 24 is a concept illustration of a total artificial
heart (TAH) embodiment of the BEEP system.
[0087] FIG. 25 a schematic view of the human torso illustrating the
location of the main components of a TAH embodiment of a BEEP
system in the human body.
[0088] FIG. 26 is a schematic view generally identifying an
alternative-component configuration of an L-VAD embodiment of a
BEEP system.
[0089] FIG. 27 is a concept illustration of the human heart
illustrating the design and location of an alternative ejection
volume measuring apparatus.
[0090] FIG. 28 is a diagrammatic illustration of the main
components of the circulation system in the BI-VAD embodiment.
[0091] FIG. 29 is a flow chart schematically illustrating the
development of the mathematical model (equations 7 and 8) for the
dynamic system including the new VAD, in this case the L-VAD.
[0092] FIG. 30 is a flow chart schematically illustrating
application of the power optimization process in a system including
the new ventricular assist device (VAD), in this case the
L-VAD.
[0093] FIG. 31 is a flow chart schematically illustrating the
multi-input, multi-output control system for performing the new
process, and the controller optimization process.
[0094] FIG. 32 is a flow chart schematically illustrating
application of the new process in a system including the new VAD in
an L-VAD arrangement.
[0095] FIG. 33 is a flow chart schematically illustrating
application of the new process in a system including the new VAD in
a BI-VAD arrangement.
[0096] FIG. 34 is a flow chart schematically illustrating
application of the new process in an alternative system including
the new total artificial heart (TAH).
DETAILED DESCRIPTION OF THE INVENTION
[0097] FIGS. 1 through 4 are schematic illustrations of the BEEP
system of the present invention, and the structural elements
thereof. For the convenience of the reader, the unique
power-optimizing and controller-optimizing methods, which are major
aspects of the invention, and they are incorporated in the new BEEP
system, are illustrated schematically by flow charts in FIGS.
28-34, to be described further, later herein.
[0098] The elements of the new BEEP system, as shown in FIG. 1, for
example, and generally designated 35, compose three primary
components: a ventricular assist device (VAD), which in this
embodiment is an L-VAD, generally designated 74 (shown on the left
side of FIG. 1). L-VAD 74 is actuated by a hydraulic pump,
generally designated 42, and controlled by a battery/controller
assembly, generally designated 65. It is to be understood that the
new BEEP system 35 will be referred to throughout this document by
the same reference numeral, in relation to a variety of
embodiments. Thus, BEEP system 35 may include a L-VAD, R-VAD,
BI-VAD, or TAH, all of which are described further herein, or the
system may include alternative embodiments of any of the VADs or
the TAH described below. The BEEP system is only a vehicle for the
other aspects of the invention, the optimization process described
in FIGS. 28-34. The optimization process can be applied to any
current or future apparatus design of L-VAD, R-VAD, BI-VAD or TAH.
The BEEP system per se, however, is nonetheless considered to be
another important aspect of the invention, regardless of which
embodiments of the various components are included in the system.
Further in regard to the various embodiments of the system, if
certain aspects of the overall system are not described in detail
as being different or distinguishable from the other embodiments,
they are considered to be the same or equivalent to those
previously or later described.
[0099] BEEP system 35 utilizes electromagnetic coils 46, 48, and 50
to drive a high ferromagnetic-constant solid cylindrical driving
magnet 40 in reciprocating fashion along the length of hydraulic
pump 42. While three such coils are preferred, it is to be
understood that the new system 35 and the alternative embodiments
thereof can operate adequately with more than or fewer than three
electromagnetic coils on pump 42. Driving magnet 40 is acting as a
piston in a hydraulic pump. The interior vessel of the hydraulic
pump may or may not incorporate end caps 56 and 57 as part of its
hydraulic-vessel design. However, the presence of the end caps made
of ferromagnetic material assist in directing the flux lines from
the surrounding coils to the driving magnet 40. It will be obvious
to those skilled in the art that several alternative embodiments
can be contemplated by changing the cross sectional areas of the
components, which may be circular, rectangular, or a number of
other closed shapes.
[0100] FIG. 5 shows pump 42 in cross section and illustrates the
external cylindrical surface of driving magnet 40 mating with the
interior cylindrical surface of electromagnetic coils 46, 48, and
50. These surfaces, whether shaped as the preferred cylinders, or
otherwise, are nevertheless sized and shaped to slidingly interact
as well as to minimize leakage of hydraulic fluid therebetween. It
is understood that the function of pump 42 is to use one or more
electromagnetic coils to drive one or more magnets in a way to
provide motion to driven magnet 44; and several alternative
embodiments can be used to accomplish this function. It will be
obvious to those with ordinary skill in the art that there can be
many variations on the cross-sectional view of the components, on
the exact orientation of the electromagnetic fields, on the exact
orientation of the magnets, on the type of hydraulic or pneumatic
fluid, and on the details of the design of the vessel containing
the hydraulic or pneumatic fluid etc, and these alternative
embodiments are herein included. It is understood that several
alternative embodiments to minimize leakage from the high to the
low pressure of the hydraulic fluid and blood, and alternative
embodiments to minimize friction between sliding components, are
conceived and considered acceptable alternative designs in the
present invention.
[0101] End caps 56, 57 are also made of ferromagnetic material, and
are disposed on opposite ends of pump 42. End caps 56, 57 are
provided with central openings 56a, 57a, so that the interior space
defined by the electromagnetic coils is in fluid communication with
main hydraulic line 60 at each end of the pump cylinder, permitting
the hydraulic fluid to flow in and out of the pump cylinder, as
described further hereafter.
[0102] End caps 56, 57 also serve to concentrate the magnetic flux
of electromagnetic coils 46, 48 and 50 in a smaller combined area,
thereby improving pump efficiency. The shape of end caps 56 and 57
assists in the optimal placement and concentration of magnetic flux
lines and minimization of the weight and dimensions of system
components. While end caps 56, 57 act as "stops" for the piston,
there may also be provided with separate "stops" of known
construction, for example, as illustrated in FIG. 19.
[0103] Magnet 40 is preferably entirely solid and thus is sometimes
referred to herein as "solid magnet 40" for convenience of the
reader. However, magnet 40 may be only substantially solid; i.e.,
there could be a small through-hole plugged with plastic, for
example, or other conceivable interruptions to the integrity of the
magnet 40 which would not prohibit system 35 from working
sufficiently in the present system. However, for most efficient,
ideal operation, magnet 40 is entirely solid.
[0104] Solid magnet 40 acts as a piston to apply force to hydraulic
fluid 52 to thereby ultimately move a driven annular magnet 44 (as
indicated by arrows A, in FIG. 1) along the length of L-VAD 74. The
magnetic flux of annular driven magnet 44 and annular valve-seat
magnet 54 are along the axial length of L-VAD 74 so that they are
biased toward and tend to lock to one another. Movement of annular
driven magnet 44 in turn moves high-ferromagnetic-constant annular
valve-seat magnet 54. Blood 78 is therefore pumped by L-VAD 74 in
the direction of the flow arrows B through aortic arch 80, as shown
in FIG. 1.
[0105] Except for hydraulic fluid leakages, the reciprocating
motion of solid driving magnet 40 is in phase with the
reciprocating motion of annular driven magnet 44, but is slightly
out of phase with the reciprocating motion of valve-seat magnet 54
due to flood, hydraulic fluid, and electromagnetic inertia effects.
The out-of-phase separation of driven magnet 44 from valve-seat
magnet 54 varies throughout the reciprocating cycle. These delays
are accounted for in the time {t} expressions in equation (7) of
the new process described later herein.
[0106] The reciprocating movement of driving magnet 40 along the
length of hydraulic pump 42 is controlled by power, voltage and
current from battery 62 to electromagnetic coils 46, 48 and 50 in
the sequence depicted in FIGS. 17, 18 and 19 and described below.
The timing and magnitude of the current from battery 62 is
controlled by controller 64 in response to ECG signals initiated
from the ECG signal 66, signals from measurements of ejected blood
volume, and as a result of the optimization process explained
below. Battery 62 and controller 64 can be connected as a
battery/controller assembly 65, as illustrated, or utilized as
separate components. The movement of driving magnet 40 is slightly
out of phase with the magnetic field along electromagnetic coils
46, 48 and 50 due to electromagnetic hysteresis effects, which are
also accounted for in the time {t} expression of equation (7)
described herein below. The out-of-phase separation of driving
magnet 40 from the magnetic field of electromagnetic coils 46, 48
and 50 varies throughout the reciprocating cycle and is
mathematically accounted for by the optimization method described
later herein to minimize the power required for operation of the
new system.
[0107] Inside an inner sleeve 68 is contained an annular valve-seat
magnet 54 which contains one-way valve 70. The sliding facing
surfaces of annular valve-seat magnet 54 and the inner sleeve 68
are sized and shaped to be substantially fluid-tight to minimize
leakage of blood therebetween. Similarly, the mating surfaces of
annular driven magnet 44 and the inner sleeve 68 and outer sleeve
72 (shown in FIG. 6) are designed to minimize leakage between
sliding facing surfaces thereof.
[0108] It is to be understood that several alternative embodiments
to minimize leakage from the various mating elements are conceived.
It is further to be understood that all elements of the new pumping
device and the entire system for operation thereof are formed of
suitable biocompatible, surgical grade materials. Such materials
may be appropriately selected from materials that are now known, as
well as new materials, which may yet be developed.
[0109] Hydraulic pump 42 drives hydraulic fluid 52 in the direction
of the flow arrows through hydraulic line 59 and into annular space
86, located between inner sleeve 68 and outer sleeve 72. The
reciprocating motion of hydraulic fluid 52 moves annular driven
magnet 44, also located between inner sleeve 68 and outer sleeve
72. By reversing the direction of current flow in electromagnetic
coils 46, 48 and 50, the direction of driving magnet 40 is
reversed, hence the direction of hydraulic fluid 52 is also
reversed; it follows that the direction of driven magnet 44 is
reversed as well. As annular driven magnet 44 is moved by the flow
of hydraulic fluid 52, magnetic interaction with valve-seat magnet
54 causes valve-seat magnet 54 to move along with annular driven
magnet 44. Because one-way heart valve, for example, as indicated
schematically at 70, is secured to valve-seat magnet 54, one-way
valve 70 moves in the same direction as valve-seat magnet 54 and
annular driven magnet 44. When one-way valve 70 moves in a
direction away from aortic valve 76, it is closed and pushes blood
78 through aortic arch 80.
[0110] One-way valve 70 can be any artificial or natural heart
valve. Some known valves are mechanical, some are biological and
some are made with compliant man-made materials. Some one-way
valves may also eventually be made with stem cell research.
Depending upon the particular type of valve selected for the
one-way valve 70, limits may be imposed on the optimization process
of equation (7), due to the pressure differences the particular
valves can withstand (e.g. prolapse may occur with some compliant
heart valves). Such differences are taken into account in the
selection for a particular system as may be necessary.
[0111] FIG. 1 depicts the state of BEEP system 35 at late systole
of the native human heart, when valve-seat magnet 54 is at the
beginning of its pumping stroke along the length of L-VAD 74. At
the stage of the cycle shown in FIG. 1 the ECG signal 66 and other
volume and pressure signals have been transmitted along wire 63 to
controller 64. (Wire 63 may also be inside of conductor 410 in the
embodiment shown in FIG. 27 and discussed hereafter.) In response
to these signals and the new optimization process, controller 64
discharges electrical power, voltage and current to hydraulic pump
42 along wires 67, 69 (and 71, in some cases). Specifically,
current from battery 62 has energized electromagnetic coils 46 and
48. In response to the energization of coils 46 and 48, driving
magnet 40 has begun to move away from its position within
electromagnetic coil 46 and has moved partially within the walls of
electromagnetic coil 48 (a cross-sectional view of driving magnet
40 and electromagnetic coil 46 is shown in FIG. 5).
[0112] Further with reference to FIG. 1, the movement of driving
magnet 40 has forced hydraulic fluid 52 to move through main
hydraulic line 60 and secondary hydraulic line 82. The motion of
hydraulic fluid 52 places pressure on both driven magnet 44 and
check valve 84. Check valve 84 is closed, as is normally the case,
securing the required pressure gradient between the high-pressure
and low-pressure imposed by the motion of driving magnet 40 within
hydraulic pump 42. Due to pressure from hydraulic fluid 52, annular
driven magnet 44 has just begun to move along the length of L-VAD
74, within annular space 86. Magnetic interactions have caused
valve-seat magnet 54 to move in a corresponding manner. Driven
magnet 44 is located slightly ahead of valve-seat magnet 54 due to
electromagnetic and fluid inertia. A cross-sectional view of L-VAD
74, through outer sleeve 72, annular driven magnet 44, inner sleeve
68, and valve-seat magnet 54 is shown in FIG. 6. The function of
the two magnets, 44 and 54, is to magnetically "lock" to each other
so that the movement of magnet 54 is affected by the movement of
magnet 44. By "lock" it is meant that the motion of one magnet
affects the motion of the other magnet via their magnetic
interaction, even though the dynamics of the system may dictate
that the motions of the two magnets may be out of phase. Hydraulic
vessel (or "sleeve" in some cases) 72 for magnet 54 and blood
vessel 68 for magnet 44 may be concentric or not, parallel or not,
and may have any cross-section. It will be obvious to those with
ordinary skill in the art that there are several alternative
embodiments for the cross-sectional view of the hydraulic and blood
vessels (parallel axes or not, concentric axes or not, circular,
rectangular or other cross section etc) and the exact location and
orientation of north and south poles of the magnets, and these are
included herein.
[0113] Aortic valve 76, located at the outlet of the left ventricle
90, has been retained open by the beginning of the movement of
one-way valve 70, which is closed and is being moved upward by
driven magnet 44. The difference in axial location of driven magnet
44 and valve-seat magnet 54 is due to fluid inertia, but also due
to magnetic inertia. Neither fluid inertia nor magnetic inertia is
accounted for in the prior art. Although in this embodiment it is
preferred that one-way valve 70 is an artificial valve of known or
newly developed variety, valve 70 may also, if desired or
necessary, be a natural heart valve or a one-way valve formed of
tissue (human or other animal).
[0114] The movement of closed one-way valve 70 is beginning to pump
blood along the length of the ascending aorta 88 and into the
aortic arch 80.
[0115] FIG. 2 depicts the state of BEEP system 35 halfway through
the pumping motion of L-VAD 74. In this figure driving magnet 40
has moved within the walls of electromagnetic coil 48,
approximately halfway through its motion along the length of
hydraulic pump 42, and electromagnetic coil 50 has been energized
by current from battery 62. The continued motion of driving magnet
40 has placed further pressure, via hydraulic fluid 52, on annular
driven magnet 44. Due to magnetic interactions with annular driven
magnet 44, valve-seat magnet 54 has moved approximately halfway
through its motion along the length of L-VAD 74. Still closed,
one-way valve 70 has pumped more blood, that would otherwise not
have been pumped by the native heart, out of the left ventricle
along the length of the ascending aorta 88 and into aortic arch 80.
Aortic valve 76 remains open, allowing the flow of blood from the
left ventricle 90 into ascending aorta 88.
[0116] FIG. 3 depicts the state of BEEP system 35 at the beginning
of the return stroke of valve-seat magnet 54. As driving magnet 40
reverses its previous motion along the length of hydraulic pump 42,
the flow of hydraulic fluid 52 through main hydraulic line 60 is
reversed as well, as indicated by the flow arrows in the Figure.
The reverse flow of hydraulic fluid 52 places pressure on annular
driven magnet 44, pushing it back along the length of the L-VAD 74,
in the direction of aortic valve 76. As annular driven magnet 44
moves back along the length of L-VAD 74, valve-seat magnet 54 and
one-way heart valve 70 move in a corresponding manner. One-way
valve 70 is open as it moves toward aortic valve 76, allowing blood
to flow freely through one-way valve 70 as it moves. Aortic valve
76 is closed at this time, preventing blood from flowing out of the
L-VAD 74 and into left ventricle 90.
[0117] FIG. 4 depicts the state of BEEP system 35 halfway through
the return stroke of valve-seat magnet 54. Driving magnet 40 has
moved back within the walls of electromagnetic coil 48,
approximately halfway through its return motion along the length of
hydraulic pump 42. The continued-motion of driving magnet 40 has
placed further pressure, via hydraulic fluid 52, on annular driven
magnet 44, pushing it back down along the length of L-VAD 74.
Valve-seat magnet 54 has moved approximately halfway through its
return motion along the length of L-VAD 74. One-way valve 70 is
still open, allowing blood to flow freely through it as it moves.
Aortic valve 76 remains closed, preventing the flow of blood from
L-VAD 74 into left ventricle 90.
Pulsatile Flow and the Present Approach
[0118] The principles of fluid dynamics require a measurable work
per cycle (and power output) from the heart to overcome the
pressure difference in the passages of the circulatory system.
Providing pulsatile instead of steady flow, accelerating and
decelerating blood and muscle, consumes significant measurable
additional work (and power) from that required for steady flow. If
the natural heart provided continuous flow under constant pressure,
then thrombi would tend to form and gradually enlarge in relatively
stagnant or low-velocity flow regions. In steady flow conditions
these thrombi would tend to become larger with time. Eventually the
larger thrombi could potentially be dislodged by the surrounding
flow causing blockage in narrower passages downstream. The results
would be disastrous. The human body would not provide more
pulsatile flow than that required for physiological reasons.
[0119] The human body requires pulsatile blood flow for survival,
and a successful artificial heart pump or VAD should emulate the
type of pulsatile blood flow provided by the native heart. Unlike
known art devices, the present invention produces an optimized
pulsatile flow. The VAD of the present invention provides the
"vector" or "matrix" difference between the unsteady flow required
by the human body and the unsteady flow provided by the native
diseased heart, hence supplying only the required deficit. By
"vector" or "matrix" difference we imply that this is not a simple
subtraction of two quantities, as it will become evident in the
following. In a total replacement configuration (TAH) the invention
provides the total unsteady flow required by the human body. While
other inventions purport to optimize the flow, the present
invention illustrates the actual requirements (engineering
principles) for this optimization.
[0120] The physical dimensions of the VAD or total replacement
heart must be optimized to each application (i.e. to each patient).
The moving mass, damping and stiffness of the combined system
(moving parts of the VAD plus native heart, if any, plus driven
blood flow through the vessels plus hydraulic fluid, surrounding
tissue, electromagnetic dynamic phenomena, etc.) must be optimized
to the dynamic response of the system (which is a form of the
natural frequencies and damping of the overall system). If these
conditions are not met, then the VAD or total replacement heart
will be inefficient; it will require more power than the minimum to
obtain the desired unsteady-flow output to the body. A good
physical example of this is a yo-yo. If the string is pulled with
the right forces at the right times (which corresponds to the
optimized forcing function for the yo-yo), it requires minimum
effort for maximum periodic travel and produces spectacular
results. If either the forcing function or the timing are not
exactly right, then it takes more effort to obtain any travel, and
the results are not as good. Another equally important aspect of
the invention is that the physical arrangement and dimensions of
the invention are optimized to the desired amplitude and frequency
of the unsteadiness in blood flow required by the circulatory
system.
[0121] Thus the power input to TAHs and VADs must be optimized to
the dynamic response of the system, otherwise the efficiency will
be low (they will require a lot of power to drive them). One of the
claims of the proposed VAD is that its driving force-time and
force-distance relationships are optimized for minimum power input
to the desired unsteady-flow characteristics, via a prescribed
procedure, thus increasing its efficiency. This is done via a
mathematical method described below. A pre-requisite for the use of
this method is a deeper understanding of details of the flow and
pressure conditions in the cardiovascular system than that in
present medical and bioengineering practice. In other words, one
needs to understand the details of the pressure and flow traces in
the native as well as the artificial systems in order to design an
efficient VAD or TAH.
[0122] While it is understood that the pressure trace changes phase
and amplitude downstream from the aorta, there is no
acknowledgement as to whether the measured pressure traces are
static, stagnation or total pressures (defined in most fluids
engineering texts). While it is clear that during most of systole
the left ventricular pressure must be higher than the aortic
pressure (otherwise flow would be in the reverse direction from the
aorta to the ventricle), some texts indicate otherwise. The premise
of this disclosure is that any TAH or VAD must be optimized around
the details of the pumping system and match the requirements of the
human body.
[0123] Static pressure p.sub.st is the pressure one would feel
while traveling along with the velocity of the fluid in a channel.
Stagnation pressure p.sub.0 is the pressure one would feel with the
fluid coming to rest against the measuring device. Total pressure
p.sub.T is the stagnation pressure plus the static head of a column
of fluid above the measuring point. For a perfect incompressible
fluid of constant density .rho. (which is one of many frequently
used mathematical models for blood) moving with velocity C the
governing equations are:
p.sub.0=p.sub.st+.rho.C.sup.2/2
p.sub.T=p.sub.st+.rho.C.sup.2/2+.rho.gz
[0124] The distinction between these three pressures in the blood
flow is important in the design of the optimal VAD, as is the
choice of measurement devices that are specialized to distinguish
measurement of static, stagnation and total pressures, and the
location of these measuring devices in the system. Optimum design
of the present device is integrally related to the fundamental laws
of fluid mechanics applied for unsteady flow conditions to the
thermodynamic system enclosing the heart and circulatory system.
Those skilled in the art of unsteady thermofluid dynamics will
recognize that the system definition is of paramount importance to
the solution of the problem, and must be defined with the accuracy
and detail suggested in the text by Gyftopoulos and Beretta (1981);
i.e. the system definition will require amounts and range of valves
for: matter, parameters or constraints, and interacting forces
between system elements. These fundamental laws are usually
expressed as one equation for conservation of mass, three equations
for conservation of momentum, for example, along (x,y,z), and a
fifth equation for the energy balance (the first law of
thermodynamics).
[0125] The following equations (1-5) are valid for any fluid
continuum (compressible or incompressible, Newtonian or
non-Newtonian), and they are general in nature. The nomenclature
used is as follows:
[0126] E.sub.t, e.sub.t=energy and energy per unit
volume,.backslash..back- slash.including internal, kinetic, and
potential energy, etc.
[0127] e=specific internal energy
[0128] {right arrow over (f)}.sub.bd=external body forcing function
per unit volume (gravity, electromagnetic, etc.)
[0129] {right arrow over (f)}.sub.sf=surface forcing function per
unit volume (resulting in stress tensor .tau.)
[0130] F.sub.nh{t}=force as a function of time from the native
Heart
[0131] F.sub.vad{t}=force as a function of time from the VAD
[0132] h=specific enthalpy
[0133] m=mass
[0134] p=pressure
[0135] Q,q.sub.x,q.sub.y,q.sub.z=heat into control volume, and heat
per unit volume in (x, y, z) coordinates
[0136] t=time
[0137] (x, y, z)=Cartesian coordinates
[0138] u, v, w=velocity components alone (x, y, z) coordinates
[0139] W=work into the control volume (from surface, from shaft,
etc.)
[0140] x=axial direction
[0141] .rho.=density
[0142] .mu.=dynamic viscosity
[0143] .tau.=stress tensor
[0144] .tau..sub.ij=element of stress tensor (includes pressure)
along (i,j) coordinates
[0145] .gradient.=divergence operator
[0146] D=total derivative operator
[0147] .differential.=partial derivative operator
[0148] The mass balance (continuity) is given by: 1 t + ( v ) = 0 t
+ ( u ) x + ( v ) y + ( w ) z = 0 ( 1 )
[0149] where the equation can be further simplified using certain
assumptions such as incompressible fluid (but here we consider the
general form of the equation with no restrictions other than
continuous fluid).
[0150] The vector form of the equation for conservation of linear
momentum can be written as the (x, y, z) momenta equations: 2 ( v )
t + ( v v ) = f sf + f bd = + f b d ( u ) t + ( u 2 ) x + ( u v ) y
+ ( u w ) z = ( xx ) x + ( xy ) y + ( xz ) z + f x , b d ( 2 ) ( v
) t + ( v u ) x + ( v 2 ) y + ( v w ) z = ( yx ) x + ( yy ) y + (
yz ) z + f y , b d ( 3 ) ( w ) t + ( w u ) x + ( w v ) y + ( w 2 )
z = ( zx ) x + ( zy ) y + ( zz ) z + f z , b d ( 4 )
[0151] where the body forces are exerted on the whole body of fluid
(such as by gravity, when {right arrow over (f)}.sub.bd=p{right
arrow over (g)}; or by external electromagnetic fields); and the
surface forces are exterted by the interior surface of the control
volume of fluid. Some texts choose to separate the pressure terms
from the stress tensor, but here the pressure terms are included in
the stress tensor .tau..
[0152] The energy balance equation is given by: 3 D E t D t = D Q D
t + D W D t ( e t ) t + ( u e t ) x + ( v e t ) y + ( w e t ) z = (
q x + u xx + v xy + w xz ) x + ( q y + u yx + v yy + w yz ) y + ( q
z + u zx + v zy + w zz ) z ( 5 )
[0153] where q.sub.x, q.sub.y, q.sub.z are the external heat
transfers (Q) in each direction, the work (W) terms are given by
tensor times velocity applied to the surface of the control volume,
and E.sub.t includes all the energy terms. For example, if these
include only internal energy, kinetic energy, and potential energy,
then E.sub.t=.rho.e.sub.t=.rho.(e+.- vertline.{right arrow over
(v)}.vertline..sup.2/2+{right arrow over (g)}.multidot.{right arrow
over (r)}). However, in the general case, E.sub.t includes all
energy terms affecting the solution of the equations.
[0154] The above five equations can be written in vector form as: 4
G t + A x + B y + C z = X x + Y y + Z z ( 6 )
[0155] where
G=[.rho., .rho.u, .rho.v, .rho.w, .rho.e.sub.t]
A=[.rho.u, .rho.u.sup.2, .rho.uv, .rho.uw, .rho.ue.sub.t]
B=[.rho.v, .rho.uv, .rho.v.sup.2, .rho.wv, .rho.ve.sub.t]
C=[.rho.w, .rho.wu, .rho.wv, .rho.w.sup.2, .rho.we.sub.t]
X=[0, .tau..sub.xx+{right arrow over (f)}.sub.x, bd, .tau..sub.xy,
.tau..sub.xz,
q.sub.x+u.tau..sub.xx+v.tau..sub.xy+w.tau..sub.xz]
Y=[0, .tau..sub.yx.tau..sub.yy+{right arrow over (f)}.sub.y, bd,
.tau..sub.xy,
q.sub.y+u.tau..sub.yx+v.tau..sub.yy+w.tau..sub.yz]
Z=[0, .tau..sub.zx, .tau..sub.yx, .tau..sub.zz+{right arrow over
(f)}.sub.z, bd,
q.sub.z+u.tau..sub.zx+v.tau..sub.zy+w.tau..sub.zz]
[0156] (the native heart and VAD forcing function are terms {right
arrow over (f)}.sub.x, bd in X, Y, Z).
[0157] The only restriction in the above equations 1-6 is that
blood behaves as a continuous fluid (if it didn't, e.g. if there is
cavitation, severe lysis, or severe clotting, then the model is
inadequate, but the resulting VAD is also useless). These equations
are valid for steady and unsteady flow (periodic and transient),
with any external force field or surface forcing function, with
heat transfer, with Newtonian or non-Newtonian fluids etc.
[0158] The resulting instantaneous equations of fluid motion (1-6)
have instantaneous eigenvalues and eigenvectors that can be
computed, and those must be matched with the combined forcing
function from the native heart F.sub.nh{t} and the VAD F.sub.nh{t},
i.e. the dynamic system of equations for the native cardio-rheology
as modified by the presence of the operating VAD. The resulting
instantaneous system of dynamic equations are of the form:
[M]{{umlaut over (x)}}+[C]{{dot over
(x)}}+[K]{x}=F{t}=F.sub.nh{t}+F.sub.v- ad{t} (7)
[0159] where [M], [C], [K] are the instantaneous non-linear mass,
damping and stiffness matrices respectively of the dynamic model.
They are non-linear because they change with time and with
mathematical or experimental data model, because the human tissue
and mechanical components are not linear, and because they change
with instantaneous position and geometry (for example with open or
closed valves), and also with daily condition of the patient. In
any case the procedures to model the dynamic system are well
established, and the fidelity of the dynamic model is improving
with time as better experimental data and theoretical or numerical
models become available for each component of the dynamic
system.
[0160] The forcing function of the native heart F.sub.nh{t} to the
dynamic system is provided by the human and can be measured (though
it also can be modeled with basic physiological interactions). The
forcing function of the VAD system is provided by the magnetic
field to the coils, which is generated by the current and voltage
to the coils, so that for a discretized dynamic system the
instantaneous power (at any time t) by the VAD is balanced
with:
W(t)=F.sub.vad{t}.multidot.{{dot over
(x)}}+losses=V{t}.multidot.i{t} (8)
[0161] where W(t) is the instantaneous power at any instant in time
t, {{dot over (x)}} are the elemental velocities at the
displacements where elemental forces F.sub.vad{t} are acting, and
the product V{t} i{t} represents the sum of the electric power
(voltage times current) supplied to the coils. In one embodiment of
the optimization procedure the physical dynamic systems are linked
so that the left-hand side of equation (7) is linked directly in
the optimization process to the right-hand side of equation (8).
The losses are electromagnetic losses of transmitting magnetic flux
from the coils to the magnets, and friction losses until this power
reaches the elemental displacements {x} on which forces
F.sub.vad{t} act, and other similar losses. These losses can be
measured or modeled mathematically with techniques available in
mechanics, fluid dynamics, electromagnetism, and other engineering
texts. Thus the model includes muscle, tissue, blood, hydraulic
fluid, and electromagnetic and mechanical effects of mass, damping
and stiffness. For example, these include friction and leakage in
the mechanical components and fluid passages, the hysteresis loop
of the electromagnetic drive of the VAD (a condition commonly
called "latching"), other electromagnetic losses, and the stress
tensors in equations 1-6, so that the resulting fluid-structure
system is driven in an optimal manner.
[0162] In general engineering systems, the power optimization and
control-scheme optimization (such as those described later for the
mechanical blood pumping device and patient) would be best applied
to the actual system itself. In this particular system, namely the
patient with the mechanical pump surgically implanted, it would be
extremely difficult to perform the steady-state power optimization
scheme, and difficult to perform the control-optimization scheme,
as this may endanger the life of the patient. These best preferred
embodiments of using the physical dynamic should eventually be
possible after clinical trials. Alternative embodiments
(alternative models) of the physical. dynamic system are likely to
be used for scientific development of the BEEP system. These are
likely to use analytic, numerical, or experimental, etc.,
expressions, or their combination, to represent the physical
dynamic system. These models can be of various degrees of
complexity. Some of these models may represent the whole dynamic
system, and others may represent only portions of the whole system.
From the above it is easy to foresee that one group of such
possible dynamic models may include the mechanical blood pump only,
while others may, in addition, incorporate portions of the native
heart and the circulatory system, etc. Similarly, one set of such
models may concentrate on finding the optimum F.sub.vad{t},
represented, for example, by forces and velocities acting on one
of: a) valve seat magnet 54; b) driven magnet 44; or c)driving
magnet 40, etc. With similar dynamic models of the electromagnetic
and hydraulic systems, this forcing function, F.sub.vad{t}, can be
correlated with the instantaneous electric power to the coils
(resulting in a form of equation (8)). The preferred embodiment of
the dynamic system model directly correlates the forces on the
left-hand side of equation (7) with the power on the right-hand
side of equation (8). Alternative embodiments of the optimization
schemes may use simplified portions of the whole dynamic system,
such as those that find the force, F.sub.vad{t} on: a) valve-seat
magnet 54; b) driven magnet 44; or c) driving magnet 40.
[0163] In this invention the forcing function F{t} in equation (7)
consists of two parts, one provided by the native heart and the
other provided by the VAD. For this purpose, "optimize" means
minimizing the power required to drive the VAD while minimizing the
shear stress imposed on the blood.
[0164] Again, the terms "optimize" and "complement" are used in
reference to the devices and systems of the present invention, it
is meant that at each heart beat and stroke of the VAD (used here
to mean either the VAD or TAH as described below), several actions
are carefully timed such that:
[0165] a) the native heart is allowed to pump as much blood as it
can on its own before the VAD is activated;
[0166] b) as the blood-ejection phase of the native heart nears
completion, the VAD is energized to provide additional pumping
action;
[0167] c) the additional pumping action reduces the back pressure
in that native ventricle so that the native ventricle pumps more
than it would have pumped unaided;
[0168] d) the timing of the action, length of pumping stroke, and
rate of pumping (stroke displacement versus time and resulting
power input versus time) of the VAD are related to the native heart
ejected blood volume and rhythm in a manner that minimizes power
input to the VAD while meeting physiological constraints;
[0169] e) the optimization processes in d) take into account the
dynamic interaction between the native heart and the VAD; and
[0170] f) the optimization process and the control scheme are
integrated with the resulting changes in blood ejected per heart
beat and heart rate (beats per minute) by the combined action of
the native heart and the VAD.
[0171] Specifically, the combination of the patient's native
cardiovascular system and the VAD at any condition of flow rate and
beating frequency supplied by the native heart will result in an
optimal shape (function of location and time as shown in the
Figure) for the forcing function provided by the VAD. The forcing
function and frequency of the VAD are controlled as explained
elsewhere in these documents. The equations presented above are
general and they are not dependent on the details of the
mathematical models. Some research teams will choose to simplify
these equations using the incompressible fluid approximation,
Newtonian fluid approximation, linear models in finite element
method programs or linearized equations in computational fluid
dynamics (CFD) approaches. All of these simplifications are fully
included in the general equations (1-7).
[0172] The L-VAD is intended to be placed between the aortic root
and the aortic arch. Thus, for VAD applications the length L and
overall outside diameter D.sub.0 of L-VAD 74 are limited by human
physiology. There is a desire to directly wrap coils 46, 48 and 50
around the length of travel of driven magnet 54, but this not
always possible, due to geometric constraints. For example, for an
adult male L.apprxeq.10 cm and D.sub.0.apprxeq.4 cm, the overall
force that can be carried by a hollow magnet 54 is a function of
the volume of the magnet, among other factors, for example, if
cylindrical, approximated by .pi.(D.sub.2.sup.2-D.sub.1.s-
up.2)l/4, where D.sub.2 and D.sub.1 are the outside and inside
diameters of the magnet and l the length, the magnetic properties
of the materials (factor k1), geometry (factor k2) and technology
of components (e.g. leakage and friction characteristics, coil
packing, heat transfer constraints (factor k3)).
f=F(k1,k2,k3)
[0173] where clearly, other factors being equal, the force is
increased by increasing D2, the outside magnet diameter. Thus the
outside diameter of electromagnetic coils used in the linear magnet
motor of prior art becomes too big for VAD to fit into the human
body in the vicinity of the aortic arch. This maximum-diameter
issue is resolved with the use of driving magnet 44, the hydraulic
fluid, and solid (or substantially solid) magnet 40. The following
table is an indication for several distinct sizes of VADs, assuming
that the diseased native heart provides 50% of the cardiac output
required by the human body:
1 Normal Output of Required Cardiac Diseased VAD Weight Height Area
Output Native Output (kg) (m) (m.sup.2) (cc) Heart (cc) (cc) Child
50 1.3 1.3 58 29 29 Teen 55 1.65 1.6 72 36 36 Avg. 55 1.75 1.7 76
38 38 Adult Female Avg. 75 1.85 2.0 90 45 45 Adult Male Large 110
175 2.3 102 51 51 Adult
[0174] As can be seen by the equations above, several standard
sizes of L-VAD 74 can be designed. As a guide, the smallest would
be a pediatric device and the largest would be for a large adult.
What follows is an example of calculations performed on a
hypothetical individual, and is not intended to in any way limit
the present invention.
[0175] The height and weight features of a person can be converted
to body surface area by using the approximate formula below. Once
body surface area is known, normal cardiac output for a given
individual can be calculated. Normal cardiac output volume per body
surface area is 45 cc/m.sup.2.
[0176] Body Surface Area (m.sup.2)=[ht (cm)].sup.0.718[wt
(kg)].sup.0.427[0.007449]
[0177] Using a 75 kg 185 cm adult male as an example, the body
surface area calculation results in a value of 2 m.sup.2. Given the
body surface area value of 2 m.sup.2 calculated above, the normal
stroke volume for the individual is 90 cc. In end-stage
cardiomyopathy, the native heart provides approximately 50% of the
required cardiac output. In the example above the native heart
would provide approximately 45 cc. Therefore, the L-VAD would have
to provide an additional 45 cc.
[0178] What follows is a general description of the approximate
sizes of an L-VAD for the above patient at one ejection volume (45
cc) and one heart rate. The procedure must be repeated several
times for different ejection volumes and heart rates before the
optimum L-VAD dimensions for the patient are decided. The procedure
is also affected by the size of available one-way valves 70,
especially if these are of the standard artificial heart valves
available commercially, which are available in several standard
diameters, usually measured in millimeters (mm).
[0179] The maximum L-VAD displacement required in this example is
45 cc. A standard 29 mm valve nomenclature is used for one-way
valve 70. This choice affects the length of L-VAD 74 as well as the
force that must drive driven magnet 44 and valve-seat magnet 54.
The axial length of driven magnet 44 and valve-seat magnet 54 is 13
mm. The wall thickness between driven magnet 44 and valve-seat
magnet 54 is 1 mm, as is the thickness of outer sleeve 72.
[0180] For certain illustrative example cases the steady state and
acceleration force required to pump blood through L-VAD 74 is 30 to
36 N (kg m sec.sup.-2). This is based on an initial estimate of the
pressure that will be supplied by L-VAD 74, multiplied by the area
of the pumping diameter. This initial estimate accounts for the
acceleration of fluids (blood and hydraulic) pumped in the system
(about 6 L in circulation), and the initial masses of the moving
components. The volume of rare earth magnet in driven magnet 44 and
valve-seat magnet 54 required to provide the 30 to 36 N is
3.83.times.10.sup.3 mm.sup.3. The resulting thickness of valve-seat
magnet 54 with a length of 13 mm is about 3.2 mm. Thus, the inside
diameter of valve-seat magnet 54 is 29 mm and the outside diameter
is 35.4 mm. The inside diameter of driven magnet 44 is 37.4 mm. The
thickness of driven magnet 44, with a length of 13 mm and an inside
diameter of 37.4 mm, must be around 2.8 mm to achieve the desired
30-36 N. Thus, the outside diameter of driven magnet 44 is 43 mm.
Therefore, the outside diameter of L-VAD 74 is 45 mm.
[0181] The stroke length required if one-way valve is to give the
required 45 cc volume is 45.7 mm. Adding this to the axial length
of valve-seat magnet (as required by the geometry of the device)
the overall axial length of the pumping portion of L-VAD 74 becomes
58.7 mm. This length will be increased to allow for the cuffs for
hydraulic fluid and blood. It is understood that several
alternative embodiments for the cross-sectional shape of the heart
valve 70, magnets 54, 44 and 40, and cuff (or "capsule") designs
for the hydraulic connections for hydraulic fluid and blood can be
used and will be apparent to one skilled in the art and thus they
are hereby incorporated in this disclosure.
[0182] The above dimensions are used to provide geometric inputs
for the models used in equation (7). The inputs result in elements
for mass matrix [M], damping matrix [C], and stiffness matrix [K].
Elements of [M] are evaluated using material densities. Elements of
[C] are evaluated using fluid dynamics for the flow passages,
structural damping for tissues and electromagnetic properties for
magnets, coils and other components, as needed. Elements of [K] are
evaluated using material and surrounding tissue properties and
electromagnetic properties for magnets, coils and other components,
as needed. The surrounding tissue must extend to the control volume
of the system where the tissue geometry is not moving. This means a
little further out of the pericardium (to fully include pericardium
tremors) and a little further out of the blood and hydraulic fluid
vessels (to include stiffness and compliance, providing elements
for [C] and for [K]).
[0183] For example, the pressure drops in hydraulic lines 60 and 82
initially can be estimated using analytic calculations available in
standard textbooks, and later evaluated by discretized mathematical
models as elements of matrices in equation (7).
[0184] Continuing the above example, with certain engineering
assumptions driving magnet 40 could be 3800 mm.sup.3 grade 37 rare
earth magnet. In one embodiment this magnet could have radius 9.7
mm, and length 12.86 mm. The hydraulic volume displaced by 45.7 mm
of stroke length of driven magnet 44 is 54.7 mm. Thus, the overall
length of hydraulic pump 42 is about 81 mm (this length will be
increased by the length of the hydraulic cuffs).
[0185] Additional secondary calculations are made to evaluate the
geometries of auxiliary components such as hydraulic lines, and
other components such as tissue in the myocardium and blood flow
system. The inputs result in elements for matrices [M], [C], and
[K] (some measured in clinical trials, others measured for
individual patients). The elements of vector of displacements {x}
and its derivatives {{dot over (x)}} and {{umlaut over (x)}} in
equation (7) are elemental displacements. The equation is nonlinear
and can be decomposed in a few or infinitely many degrees of
freedom, depending on the fidelity of the dynamic model.
[0186] F.sub.nh{t} in equation (7) is measured for each condition
(heart rate, ECG signal, volumes ejected from right and left
ventricles, and pressures) of the patient. For L-VAD 74 this is
given by the total pressure (static+dynamic+elevation components)
provided by the diseased heart inside the left ventricle as a
function of time (measured during the heartbeat) integrated over
the inside surface area of the four chambers of the heart. This
surface area is also measured with magnetic resonance imaging
(MRI), echocardiography, or other similar technique. These give
pressure-volume-time traces for the diseased heart as illustrated
in FIGS. 14 and 15. The volume information is correlated with data
from proximity sensors, such as 406 and 416 in FIG. 9, which may
be, for example, proximity sensors. This information changes as the
condition of the patient worsens or improves. This means that the
data needs to be calibrated before surgery, and again soon after
surgery, and monitored periodically so that the data provided by
proximity sensors 406 and 416 reflect the forcing function provided
by the native heart, where the mathematical expressions are:
dF.sub.nh{t}=p(t)dA(t)
F.sub.nh{t}=.intg..sub.Ap(t)dA(t)
[0187] The resulting pressure-volume-time traces of the native
heart have a phase associated with the timing of the forcing
function F.sub.nh{t} during the beat. This can be modeled with
Fourier series analysis of the pressure and volume signals of the
native heart over time. Again, these vary with the rate (beats per
minute) of the native heart and with the condition of the patient
(i.e. the information changes as a function of time and needs
periodic updating).
[0188] FIG. 6 is a cross-sectional view of L-VAD 74, along line 6-6
of FIG. 1. By contrast, FIG. 7 is a concept illustration of the
human heart showing the location of L-VAD 74 in place of ascending
aorta 88. It is understood that the L-Vad may replace a portion,
not necessarily the entire ascending aorta. Further in some
embodiments the aorta may simply be transected to place the L-VAD
outside the body, with blood conduits connecting the ends of the
transected aorta to the L-VAD. Corresponding alternative
embodiments are possible for the R-VAD, BI-VAd. Blood moves from
left ventricle 90, through aortic valve 76 and into L-VAD 74,
situated in place of the ascending aorta 88, pumps blood into
aortic arch 80.
[0189] FIG. 8 shows the placement of an entire BEEP system 35
within the human torso. The illustration depicts the spatial
relationship between battery/controller assembly 65 and L-VAD 74.
FIGS. 7 to 13 and 20 to 25 are schematic illustrations, not
cross-sectional views, and the location of L-VAD 74 in FIG. 8 is at
a different plane from the location of R-VAD 58 in FIG. 21. (The
L-VAD in FIG. 8 is correctly shown more to the right of the
patient's chest than RVAD in FIG. 21, but FIGS. 8 and 21 are
anatomically correct, while FIGS. 7 and 20 are simple arrangement
illustrations).
[0190] FIG. 9 is a concept illustration of the human heart,
illustrating a technique to measure the volumes of the left and
right ventricles, which is used in the control algorithm. Rare
earth (or similar material) magnets 402 and 404 are embedded in the
endocardial surface of the right ventricle, and their relative
motion changes the magnetic field between them. These changes are
measured by proximity sensor 406, mounted on the pericardium. The
signal is transferred by electrical lead 408 to wire bundle 410.
Rare earth or similar material magnets 412 and 414 are embedded in
the endocardial surface of the left ventricle, and their relative
motion changes the magnetic field between them. These changes are
measured by proximity sensor 416, mounted on the pericardium. The
signal is transferred by electrical lead 418 to wire bundle 410.
The signals from wire bundle 410 are transmitted to controller 64
and used as described later.
[0191] The proximity sensors are currently available devices that
may operate on the resistive, capacitative or inductive principles,
or combinations, or other similar distance-measuring technology.
Auxiliary (parallel horizontal) lines 1 through 6 in FIGS. 10
through 13 represent the motion of the ventricles and the KG
diaphragm (see below) during a cardiac cycle. It has traditionally
been thought that the valves of the heart open to let the blood
through when the chambers contract, and snap shut to prevent it
from flowing backward as the chambers relax. While this is correct,
the valves also act as pumping pistons for at least a portion of
the cardiac cycle, a fact not known to be previously recognized in
the literature. The plane of the valves and the supporting tissue
on the perimeter of the valves form an internal diaphragm,
approximately in the horizontal plane, which buckles and moves in
3D, which is also not known to be named in the existing literature.
For the purposes of this document this diaphragm will be the
Korakianitis-Grandia (KG) diaphragm, illustrated in FIGS. 10 to 13,
and generally designated at 92. KG diaphragm 92 has four quadrants
with a valve in each quadrant. It is activated by the surrounding
cardiac muscle, which forces the diaphragm into a
periodically-changing three-dimensional surface. (Thus FIGS. 10-13
are illustrations rather than cross-sections of the human heart).
During the cardiac cycle the aortic valve and pulmonary valve stay
nearly immobile (which allows one to place the VAD on the outlet
side of these two valves); but the mitral and tricuspid valves move
substantially, contributing at least in part to the pumping action
of the ventricles. The mitral valve movement is comparable to the
movement of the inside wall of the left ventricle. The tricuspid
valve exhibits an even greater excursion and corresponding pumping
action and is actually used in current medical practice as a
measure of right ventricular ejection fraction.
[0192] While the exterior surface of the heart moves slightly
during the cardiac cycle, the volume of the four-chamber heart does
not change appreciably in time. However, the known art does not
recognize that the total overall volume of each of the two sides,
left and right, of the heart does not change appreciably during the
cardiac cycle, even though the ventricular and atrial septa move.
In operation of the native heart, during left ventricular ejection,
the left atrium concurrently expands (while filling for the next
cycle), and KG diaphragm 92 begins to move towards the apex of the
heart, with complementing motions of the atrial septum and of the
ventricular septum, thus keeping the overall volume of the left
side of the heart about constant. Apex of the heart is a common
term for the tip of the left ventricle. Similar arguments keep the
right-side volume approximately constant, while the right and left
sides of the heart expel blood to the lungs and the aorta about
simultaneously. Minor deviations from these equal-volume
considerations on each side, right and left, occur due to one side
of the heart beating slightly before the other, heart-muscle and
blood-vessel elasticity, transient accelerations or decelerations
of the overall cardiac cycle (governed by the body's demand), and
blood compressibility (which at circulatory system pressures is
practically negligible).
[0193] The human cardiac cycle consists of two phases,
conventionally called diastole and systole. During diastole (FIGS.
10 and 13) the ventricular muscle is relaxing, KG diaphragm 92
moves toward the base of the heart while the aortic and pulmonary
valves are closed, and the mitral and tricuspid valves are open and
moving towards the base of the heart, thus increasing the volume
inside the ventricles while concurrently decreasing the volumes
inside the atria. Base of the heart is a common term for the
posterior aspect of the heart, behind the atria in the heart's
anatomical position. The open mitral and tricuspid valves move
upwardly (when the body is upright) to engulf blood from what was
volume inside the atria (thus concurrently increasing the volume
inside the ventricles while decreasing the volumes inside the
atria). In a model of ventricular flow this volume exchange would
affect the thermodynamic system definition, mentioned previously. A
scrutinizing review of echocardiography tapes reveals that radial
volume changes around the vertical axes of the ventricles account
for roughly 75% of the volume change, with the corresponding
movement of the KG diaphragm accounting for the remaining 25%
volume change. There is heart-muscle work associated with these
changes in volume that must be accounted with the correct
mathematical model in any attempt to model flow in the heart or in
VAD mimicking the function of the heart.
[0194] By the end of diastole, the relaxing ventricle allows KG
diaphragm 92 to move toward the base of the heart. During initial
systole (FIG. 11) the aortic and pulmonary valves remain closed
while pressure is building up inside the ventricles. Subsequently
the blood pressure inside the left ventricle becomes higher than
the pressure in the ascending aorta, and the aortic valve is opened
by blood flowing out of the ventricle. Substantially concurrently
the blood pressure inside the right ventricle becomes higher than
the pressure in the pulmonary trunk, and the pulmonary valve opens.
The motion of the KG diaphragm carrying the mitral valve toward the
apex of the heart, along with the simultaneous concentric
contraction of the ventricle, ejects blood into the ascending
aorta.
[0195] Correspondingly, the movement of KG diaphragm 92 carrying
the tricuspid valve towards the apex of the heart, along with the
simultaneous contraction of the right ventricle, ejects blood into
the pulmonary trunk. This same motion of the KG diaphragm with the
tricuspid and mitral valves closed increases the volumes inside the
atria, hence refilling the atria with blood from the pulmonary
veins (left atrium) and the vena cava (right atrium). Towards the
end of systole the aortic and pulmonic valves close, then the
mitral and tricuspid valves open and the cycle starts anew. Each
cycle takes approximately one second.
[0196] The double throb ("lub dub") of the beating heart is
generated by the snapping of the closing valves, but also from the
accompanying vibrations of the surrounding heart muscle and
contained blood.
[0197] The three-dimensional motion of KG diaphragm 92 forces each
one of the four one-way valves to act as pumping pistons for at
least part of the cardiac cycle. The blood flow must be optimized
around artificial heart valves to provide the desirable flow and
pressure pattern while minimizing shear stresses on blood
cells.
[0198] FIG. 10 represents late diastole. At this point KG diaphragm
92 is at its uppermost position (horizontal lines 1-3). Pulmonary
valve 94 and aortic valve 76 are closed, and mitral valve 96 and
tricuspid valve 98 are open, completing filling of ventricles 106
and 90 following contraction of right atrium 100 and left atrium
102. Ventricular myocardium 104 is in its relaxed state.
[0199] FIG. 11 represents early systole. At this point ventricular
myocardium 104 is thickened concentrically, and KG diaphragm 92 is
moving downward (horizontal lines 2-4), the tricuspid valve side
more so than the mitral valve side. Due to this motion, the volume
of ventricles 106 and 90 is decreased while that of atria 100 and
102 is increased. Hence, the total volume of the heart remains
essentially constant.
[0200] FIG. 12 represents late systole. At this point ventricles
106 and 90 have maximally thickened concentrically, and KG
diaphragm 92 has been pulled maximally downward (lines 4-6). This
completes the emptying of the ventricles.
[0201] FIG. 13 represents early systole. At this point KG diaphragm
92 is beginning to return upward (lines 3-5) toward atria 100 and
102, while ventricular myocardium 104 relaxes concentrically. As a
result of this motion, the volume of atria 100 and 102 decreases
while that of ventricles 106 and 90 increases, hence the overall
volume of the heart remains essentially constant.
[0202] There are minor variations to the basic steps outlined
above, due to damping and elasticity of the heart tissues, and
small amounts of native heart valve leakage, which can be accounted
for in the thermodynamic system definition mentioned earlier.
[0203] FIG. 14 represents pressure-volume loops for healthy (solid
lines) and diseased (broken lines) hearts, with pressure plotted
along the Y axis and volume plotted along the X axis. Line 112
represents a healthy heart. Point 118 to point 120 represents
ventricular filling. Point 120 to point 122 represents
"isovolumetric contraction". Point 122 to point 124 represents
ejection during systole. Point 124 to point 118 represents
isovolumetric relaxation.
[0204] Line 114 represents the normal response in accordance with
the Frank Starling law to an increase in volume. Points 126, 128,
130 and 132 correspond to points 118, 120, 122, and 124, on line
112, respectively, and represent the corresponding phases of the
cardiac cycle. Notice that points 124 and 132 lie on the same line,
commonly referred to as the End Systolic Pressure Volume
Relationship (ESPVR).
[0205] Line 116 represents a diseased heart which has exceeded the
limits of the Frank Starling curve. In these hearts, the end
diastolic pressure and volume are elevated but the end systolic
pressure is decreased from that associated with normal myocardium,
as noted by a lesser slope of the ESPVR line. Points 134, 136, 138
and 140-correspond with points 118, 120, 122 and 124 of line 112,
respectively.
[0206] FIG. 15 represents ventricular pressure over time of the
healthy and diseased hearts. Again, the end diastolic ventricular
pressure is greater in the diseased heart than in the healthy
heart, and because the ejection fraction is decreased in the
diseased heart, the heart rate is increased so that the total
cardiac output is maintained.
[0207] VADs are activated by either ECG signal, or via a
fill-to-empty mode. The control algorithm of the present device
utilizes inputs from the ECG of the native heart, as well as a
measurement of the ejection volume and pressures of both the right
and left ventricles. As a result, the prior art complication of
mismatch of ejection volume between the right and left ventricles
is eliminated.
[0208] FIG. 16 is a concept illustration of the pumping travel of
new BEEP system 35, as compared to that of known cardiac pumping
devices. The abscissa is the number of beats per minute of the
native heart. The ordinate is the length of travel of driving
magnet 40 along the length of hydraulic pump 42. The known devices
are either on or off (travel X is always equal to maximum travel
Xmax), as shown by the dotted line. In the present device, the
length of travel of driving magnet 40 along hydraulic pump 42
varies, following solid line 142, depending on the number of beats
per minute of the native heart. As the number of beats per minute
of the native heart increase, they reach high threshold dotted line
144 at which point controller 64 signals for driving magnet 40 to
start moving along the length of hydraulic pump 42 smoothly
increasing the stroke travel, approaching solid line 142.
[0209] Over time, due to augmentation of ejected volume by the VAD,
the end diastolic volume of the native heart decreases. This allows
the internal volume of the ventricle to become smaller, which
subsequently allows the muscle of the native heart to begin to
recover, and hence eject a greater volume of blood per stroke. As
cardiac output is the product of ejected blood volume times heart
rate, the increased ejected volume allows the heart rate to
decrease. These changes are sensed by controller 64, which reduces
the stroke length of the VAD as a greater portion of the cardiac
output is now being supplied by the native heart. The stroke length
of the VAD progresses to the left, from point 148 toward point 146.
When the beats per minute reach the lower threshold point 146
(reflecting at least partial recovery of the native myocardium),
this will cause a decrease in stroke length along line 150, which
effectively reduces the stroke length of the VAD. If recovery of
the native heart continues, the stroke volume of the L-VAD is
reduced to zero along line 150. At this point the native heart is
once again providing the total cardiac output on its own, without
the assistance of the VAD. The shape of line 142 is actively
manipulated by controller 64 using additional inputs for the
measurement of the ejection volume and pressures of the right and
left ventricles provided by the measurement system shown in FIG. 9
and described in relation thereto.
[0210] A critical factor in the success of a newly-installed L-VAD
is satisfactory operation of the right ventricle during the
immediate peri-operative period. The ejection fractions of the
right and left ventricles are monitored by mechanisms such as those
illustrated in FIG. 9 and alternative embodiments thereof. This
allows the volumetric outputs of the right ventricle and the
assisted left ventricle to be the same by manipulating the shape of
line 142 up or down. For example, suppose that a short time after
activation, the right ventricle is measured to give 50 cc per beat
and the left ventricle gives 25 cc per beat. In this case, the
stroke length of L-VAD 74 will be adjusted to give 25 cc per beat.
If, a short time later, the right ventricle starts to fail and now
only ejects 40 cc per beat (while the left ventricle still gives 25
cc per beat), this discrepancy in the ejection fractions will be
detected by proximity sensors 406 and 416. In response, controller
64 will lower the level of line 142, resulting in a shorter stroke
length of L-VAD 74, to give 15 cc per beat, for a total of 40 cc
from the assisted left ventricle. The human body will compensate
with a corresponding increase in heart rate (beats per minute). In
this fashion, controller 64 matches the ejection volumes of the
right and left sides. Should the right ventricle continue to fail,
a decision will have to be made as to the appropriateness of
installing an R-VAD, making this a BI-VAD system.
[0211] Assuming that a patient is supported on an L-VAD alone, the
length of the pumping stroke of the L-VAD is determined as a
function of beats per minute, as shown in FIG. 16, and manipulated
by matching the ejection volumes of the left and right sides of the
heart. For example, starting from point 148, if the beats per
minute continue to increase, then the piston stroke also continues
to increase smoothly to point 178. Starting from any operating
point to the right of point 178, as beats per minute decrease, at
point 178 the device reduces the travel of driving magnet 40 from
its maximum travel along hydraulic pump 42. If the number of beats
of the native heart is reduced sufficiently to reach low threshold
line 150, then the travel of driving magnet 40 is reduced smoothly
until it becomes zero following low threshold line 150. Lines 150
and 142 may coincide over the length of line 150. Line 142 to the
left of point 146 in FIG. 16 can represent the initial activation
of the L-VAD after surgical installation thereof.
[0212] The locations, magnitudes and exact shape of lines 142, 150
and 144 shown in FIG. 16 are for purposes of illustration and will
vary from patient to patient, and device to device. In addition,
during normal operation of the L-VAD, small up or down variations
of the level of line 142 are made by controller 64 in order to
match the ejection volumes, measured as described in FIG. 9, of the
left and right sides of the system. The control algorithm has
several input variables; among others, beats per minute measured by
the ECG (as described below), the ejection volume and pressures
from the right and left side of the -system as illustrated in FIG.
9. For clarity, in the remaining Figures the concepts are
illustrated using beats per minute to represent the function of
controller 64.
[0213] FIG. 17 is a chart comparing the activation sequence of the
known Larson et al. device and that of the present BEEP system, as
well as the corresponding power requirements, in relation to the
ECG trace of the native heart. The solid lines represent the
present device and the dotted liens represent the known device of
Larson et al. The common abscissa is the time period required for
two beats of the native heart. The ECG trace has characteristic
spikes Q, R, S (commonly known as the QRS complex), and waves T and
P, whose physiological function and importance is described in
detail in medical texts. The beginning and end of the stroke of
Larson's device occurs at or near point R of the QRS signal.
[0214] The graph shown in portion A of FIG. 17 shows
non-dimensional stroke distance traveled by driving magnet 40 (X to
Xmax) from 0.0 to 1.0, according to FIG. 16. Portion B of FIG. 17
illustrates a typical ECG voltage trace during cardiac operation.
The beginning of the stroke of the present system is at or about
the beginning of the T wave, allowing the rapid ejection phase of
the native heart to precede the augmentation of the VAD. The
pumping phase (points 154 to 156) of the present system occurs
between the end of the T wave and the beginning of the P wave. The
return stroke begins at this time (point 156) and ends at or just
after the QRS complex (point 158).
[0215] Driving magnet 40 rests at the center of electromagnetic
coil 46 (X=0) during the time period between the end of the return
stroke (point 158) and the beginning of a new stroke (point 160).
The resting period between the end of the return stroke (point 158)
and the beginning of the next stroke (point 160 corresponding to
154) is important for a number of reasons. For example,
acceleration at the beginning 154 and end 156 points of the stroke
is minimal. This resting period allows time for depolarization of
electromagnetic coils 46, 48 and 50 between strokes. When
necessary, it also allows driving magnet 40 to be centered within
electromagnetic coil 46, thereby allowing driven magnet 44 and
valve seat magnet 54 to return to the beginning of the stroke of
L-VAD 74. This return function is accomplished by opening and
closing check valve 84, as necessary.
[0216] Due to leakage of hydraulic fluid around magnets 40 and 44
it is possible that one of the two magnets is stopped at one of the
two ends of its travel while the other magnet is somewhere in the
middle of its stroke. For example, if annular driven magnet 44 is
at the end of its travel at the pump inlet (by the aortic valve as
shown in FIG. 1) but driving magnet 40 is not yet all the way back
to the beginning of its stroke (as shown in FIG. 4), then high
hydraulic pressure will arise between driving magnet 40 and end cap
57. This condition would be sensed by the large increase in the
power required by the coils. At that time the controller would open
check valve 84 and would pull driving magnet 40 by the coils
towards end cap 57, until it touches end cap 57, thus bringing the
two magnets back into phase, and normal operation would resume. The
procedure is similar if driving magnet 40 reaches the end of its
pumping travel (as shown in FIG. 3) while driven magnet 44 is near
the middle of its travel (as shown in FIG. 2). It is also possible
to correct for these leakages at the end of every pumping stroke,
or ever few pumping strokes. A similar procedure can be used for
initial activation of the device, to start the device after it has
been stopped, and to re-lock magnets 54 and 44 if they are not
locked relative to each other at any time during operation. The
latter condition is sensed by a large decrease in the power
required by the coils.
[0217] The shape of line 164 in FIG. 17 (portion A) is determined
by the optimization procedure described later herein. Acceleration
begins smoothly (point 154 on line 164) so that less power is
required than if the device started with a constant velocity.
Maximum acceleration is achieved somewhere in the middle of the
stroke, based on the optimization procedure. The, VAD of the
present system approaches maximum stroke travel with minimum
velocity at point 156 so that it does not impact against the
mechanical stop at the end of travel and no energy is lost due to
impact. Thus, less energy is required to start the return stroke.
The return stroke is less critical than the pumping stroke because
one-way valve 70 is open and less energy is required to return the
new VAD (e.g. L-VAD 74) to its starting position. Even so, the
shape of line 162 is optimized by the procedure. Velocity is zero
at X=Xmax, requiring less power than if there was a change in
velocity at Xmax. The shape of lines 164 and 162, and resting
period between points 158 and 160 is optimized by the
procedure.
[0218] FIG. 17C shows the power requirements of the present device
in Watts during usage. The maximum power requirement (point 166)
occurs somewhere along line 164, as optimized by the procedure.
During the return stroke, power peak at point 168 occurs slightly
before X=Xmax at point 156, which corresponds to the power
requirement at point 170. This occurs because of the sequence of
energization of electromagnetic coils 46, 48, and 50 as explained
further in FIG. 18.
[0219] In comparison, the known device, illustrated by the dotted
lines in FIG. 17, begins the pumping stroke at or near R of the QRS
complex, with constant velocity, until the point of maximum travel,
which occurs at or near the end of the T wave. The return stroke is
with constant velocity from the point of maximum travel until the R
peak of the next QRS complex, requiring large acceleration at the
two ends 172 and 174 of its stroke. This requires correspondingly
large power input. In addition, the velocities are not optimized
for the unsteady flow and the time varying magnetic fluxes,
requiring large power input at all points during the stroke. Points
172 and 174 correspond to power peaks 176 and 178, respectively. As
a result, the present device will take less power (solid line 180)
than the prior art device (dotted line 182), and the peaks occur at
different times.
[0220] Further with reference to FIG. 17, power peak 166 of the
present device occurs during the T wave of the native heart,
allowing the native heart to finish its rapid ejection. This
increases the volume of blood pumped due to the combination of the
native heart and the VAD with respect to the prior art (due to
summation of volume), requires less power than the prior art
device, and allows the native heart a chance to recover by
decreasing left ventricular volume. These combinations make the
BEEP system bio-compatible.
[0221] FIG. 18 is a schematic illustration of the embodiment of
BEEP system 35 shown in FIG. 1, using three electromagnetic coils
46, 48 and 50 along the length of hydraulic pump 42 showing the
corresponding magnetic flux of the driving magnet and the
electromagnetic coils. Vector 184 represents the magnetic flux of
driving magnet 40. Position X=0 is at the center of electromagnetic
coil 46. Position X/Xmax is at the center of electromagnetic coil
50. The top portion of the Figure shows electromagnetic coils 46,
48 and 50 and positive stops 186 and 188. The middle portion of the
Figure is an illustration of the typical periodic representation of
the magnetic fluxes of electromagnetic coils 46, 48 and 50 during
the pumping stroke from X=0 to X=Xmax. The bottom portion of the
Figure is an illustration of the typical periodic representation of
the magnetic fluxes of electromagnetic coils 46, 48 and 50 during
the return stroke from X=Xmax to X=0.
[0222] With reference to the middle portion of FIG. 18, the
abscissa is the axial position of driving magnet 40 from the center
of coil 46 (point 190) to the center of electromagnetic coil 50 at
X=Xmax (point 192).
[0223] With reference to the bottom portion of FIG. 18, the
abscissa is axial position of driving magnet 40 from the center of
coil 50 at X=Xmax (point 192) to the center of electromagnetic coil
46 at X=0 (point 190). Starting from X=0 in FIG. 18(b) (point 190,
which corresponds to point 154 in FIG. 17), positive stop 186
ensures that activation of magnetic fields 194 and 196 in
electromagnetic coils 46 and 48, respectively, will force driving
magnet 40 from the center of electromagnetic coil 46 towards the
center of electromagnetic coil 48. Magnetic field 194 is reduced to
zero soon after driving magnet 40 is a little outside
electromagnetic coil 46. As driving magnet 40 approaches the center
of coil 48, magnetic field 196 in coil 48 is reversed in direction
to magnetic field 198. The reversal in magnetic field 196 does not
necessarily coincide with the point in time when magnet 40 is at
the center of coil 48. The exact location of reversal is dependent
upon an optimization procedure.
[0224] Magnetic field 200 is initiated just before driving magnet
40 enters electromagnetic coil 50. The combination of magnetic
fields 198 and 200 die out by point 192 (corresponding to point 156
in FIG. 17) and smoothly bring the magnet to position X=Xmax,
against positive stop 188. At that position, the magnetic fields in
coils 48 and 50 are reversed as shown at point 192 in FIG. 18(c).
Positive stop 188 ensures that magnetic fields 202 and 204 from
coils 50 and 48, respectively, push driving magnet 40 from X=Xmax
(point 192) towards the center of coil 48. Magnetic field 202 is
reduced to zero soon after driving magnet 40 is a little outside
electromagnetic coil 50. As driving magnet 40 approaches the center
of coil 48, magnetic field 204 in coil 48 is reversed in direction
to magnetic field 206. The reversal in magnetic field 204 does not
necessarily coincide with the center of coil 48.
[0225] The exact location of reversal is dependent upon the
optimization procedure. Magnetic field 208 is initiated just before
driving magnet 40 enters electromagnetic coil 46. The combination
of magnetic fields 206 and 208 die out by point 190 (corresponding
to point 158 in FIG. 17) and smoothly bring the magnet to position
X=0, against positive stop 186. At that position, from point 158 to
point 160 in FIG. 17, magnetic field 208 (or its residual effects)
will retain driving magnet 40 at X=0, whereupon the cycle repeats
itself. All of the magnetic fields 194-208 will be optimized to
give F.sub.vad{t}. The power to obtain magnetic fields 196-208 will
be minimized based on the resistance, inductance, and capacitance
of the electromagnetic system, including the coils and magnets, and
voltage source, using constitutive relations or experimental data
for the dynamic representation of the systems in equation (7),
established in electromagnetic theory.
[0226] In general, the magnitudes of magnetic fields 194-200 will
be greater than those of magnetic fields 202-208, because the
former occur during the pumping phase with one-way valve 70 closed
and pushing blood, while the latter occur during the return stroke
with one-way valve 70 open. Even though the above embodiment
utilizes three electromagnetic coils, it is contemplated that the
present device may contain more or fewer electromagnetic coils. In
the limit, a linear stepper motor may be used.
[0227] FIG. 19 is a schematic illustration of the embodiment of
Beep System 35 illustrated in FIG. 1, wherein only two
electromagnetic coils 46 and 48 are activated and used to move
driving magnet 40. This is done in order to obtain some stroke
length X less than Xmax, as shown in FIG. 16. Similar relative
functioning, as described with respect to FIG. 18, is found in the
displacements and magnetic fluxes of FIG. 19 as well.
[0228] FIG. 20 is a concept illustration of the human heart
depicting the placement of R-VAD 58 in place of a portion of the
pulmonary trunk 210. Blood moves from right ventricle 106 through
pulmonary valve 94 and into the pulmonary trunk 210 before its
bifurcation point 212.
[0229] FIG. 21 shows the placement of an R-VAD embodiment of BEEP
system 35 within the human torso O. The illustration depicts the
spatial relationship between battery battery/controller assembly
65, hydraulic pump 42, and R-VAD 58. As mentioned previously, FIGS.
7, 9-13 and 20, 22, 24 and 27 are simple arrangement illustrations,
not anatomically-correct views.
[0230] FIG. 22 shows the placement of a BI-VAD, generally
designated 77, which consists of a combined assembly of L-VAD 74
and R-VAD 58 in the same system. In this embodiment L-VAD 74 is
located in place of at least part of the ascending aorta 88. In use
of BI-VAD 77 blood moves from left ventricle 90 through the aortic
valve 76 and into ascending aorta 88; i.e. in this system L-VAD 74
pumps blood into the aortic arch 80, just as in use of the L-VAD
alone. RVAD 58, as part of the BI-VAD 77, is located in place of at
least part of the pulmonary trunk 210, just as it is used in the
embodiment (R-VAD alone) shown in FIG. 20. Blood moves from the
right ventricle 106 through pulmonary valve 94 and into pulmonary
trunk 210 as R-VAD 58 portion of BI-VAD 77 pumps blood into the
bifurcation 212 (hidden from view) of the pulmonary arteries.
[0231] FIG. 23 shows the general placement of the BI-VAD 77
embodiment of the BEEP system in the human torso O. The
illustration depicts the relative spatial relationship of BI-VAD 77
in the chest and of battery/controller assembly 65 and hydraulic
pump 42 in the abdomen. (The L-VAD in FIG. 23 is correctly shown
more to the right of the patient's chest than RVAD, but FIG. 23 is
anatomically correct, while FIG. 22 is a simple arrangement
illustration).
[0232] FIG. 24 is a schematic illustration of the Total Artificial
Heart (TAH) embodiment, generally designated 95, for use in a
variation of the new BEEP system 35. In this TAH embodiment, atria
100 and 102, along with ECG signal 66, are retained from the native
heart. The TAH is comprised of a BI-VAD system with a greater
stroke volume than the assist embodiment, as the total cardiac
output is now being supplied by the BEEP system. In some TAH
embodiments it is possible to use the mitral valve 96 and tricuspid
valve 98, shown in previous Figures, provided their papillary
muscles and chordae tendineae are functioning, and insert he L-VAD
and/or R-VAD portions of the BI-VAD into the respective ventricles.
However, in other alternative TAH embodiments artificial valves may
be necessary or preferred. In addition, all, some or non of the
native ventricle may be retained. In the embodiment illustrated in
FIG. 24 L-VAD 74 completely takes the place of left ventricle 90
(seen in FIG. 7, for example), and hence its inlet is grafted to an
artificial valve (not shown) and its outlet is grafted into the
ascending aorta 88. Right ventricle 106 is replaced with an R-VAD
58, which has its inlet grafted to an artificial heart valve (not
shown) and its outlet grafted into the pulmonary trunk 210.
[0233] FIG. 25 shows the general placement of the TAH embodiment 95
of the new BEEP system within a human torso O. The illustration
depicts the spatial relationship among battery 62, controller 64,
in the abdomen, and TAH 95, in the chest cavity. In various
embodiments of the TAH a portion or all of the native heart may be
removed. If the chordae tendinae and papillary muscles are intact,
then the TAH would consist of an L-VAD and an RVAD placed inside
the respective ventricles. If the mitral and tricuspid valves of
the native heart are not utilized, then the TAH would require a
one-way valve at the inlet of the L-VAD and a one-way valve at the
inlet of the RVAD, and the combination that would make the overall
TAH. If the entire native heart (including the sino-atrial node) is
removed, then the L-VAD and RVAD system that would comprise the TAH
will be triggered by an electrical signal driven by sensors that
indicate the level of oxygen in the blood stream and other sensors
of body functions.
[0234] FIG. 26 represents an alternative embodiment, generally
designated 35', of an L-VAD version of BEEP system 35. In this
version, two or more electromagnetic coils, 300, 302 and 304 are
used to drive valve seat magnet 54 in a reciprocal fashion. In this
alternative embodiment, no hydraulic pump is required. The Figure
represents the beginning of the pumping stroke, at which time
one-way valve 70 is closed and electromagnetic coils 300, 302 and
304 are energized in a manner similar to that illustrated in FIGS.
18b or 19b to drive valve seat magnet 54 down the length of L-VAD
306, pumping blood into aortic arch 80. At the end of the pumping
phase, the direction of current in electromagnetic coils 300, 302
and 304 is changed, again in a manner similar to that depicted in
FIGS. 18c or 19c, driving valve seat magnet 54 (this time with
one-way valve 70 open) back down the length of L-VAD 306 to its
original position. Electromagnetic coils 300, 302 and 304 are
energized by controller 64, in response to the ECG signal 66, and
the ejection volumes and pressures of the right and left sides of
the system as illustrated in FIG. 9, in a similar manner as
described in the preferred embodiment.
[0235] FIG. 27 represents an alternative embodiment of the ejection
volume measuring apparatus for use in an alternative BEEP system,
generally designated 35" (only a portion of which is shown). In
embodiment 35", instead of rare earth magnets 402, 404, 412, and
414 (shown in FIG. 9), any proximity sensor 502, 504, 520 and 522,
can be substituted. Such proximity sensors may be included in any
or all of the four chambers of the heart. These proximity sensors
may or may not need to be coupled with conductors 506, 508, 524,
and 526. Proximity sensor output can be fed to signal converters
512 and 530 directly, or with conductors 510 and 528. The output
from signal converter 512 is directed with lead 408 into wire
bundle 410, and the output from signal converter 530 is directed
with lead 418, also into wire bundle 410. Wire bundle 410 can, but
does not necessarily, include inside it conductor 63, shown in
FIGS. 1-4, which transmits the ECG signal to controller 64. The
signals from these conductors become inputs to controller 64.
[0236] Optimization Procedure and Control Sequence:
[0237] The following optimization process can be applied to any VAD
or TAH device. During normal operation the native heart interacts
with the VAD so that the overall system/control responds to X/Xmax,
the ECG signal, and .phi. of the combined system of the native
heart and VAD.
[0238] FIGS. 28-34 illustrate the basic procedure of the
optimization and control process for new BEEP system 35, 35' or
35". It is to be understood throughout the discussion herein that,
unless otherwise specified, "VAD" shall be interpreted to mean
either the L-VAD, R-VAD, BI-VAD or TAH.
[0239] The main steps of the process are the following:
[0240] Develop a mathematical model of the dynamic behavior of the
desired part of the physical system
[0241] Identify the system inputs, outputs and desired constraints
for the physical part of the system in (a) above
[0242] Optimize power input to the VAD to complement the action of
the diseased native heart
[0243] Develop an optimized control scheme for the inputs and
outputs
[0244] Perform tests on the individual patient
[0245] Maintenance: updating the dynamic optimization and control
schemes as the condition of the patient changes with time.
[0246] The inlet and outlet boundary conditions, and other
engineering inputs to the model, and corresponding dynamic inputs,
will depend on the dynamic system definition (Gyftopoulos and
Beretta, 1991). Several alternative embodiments of the definition
of the dynamic system that will perform the optimization can be
defined, and the following examples are given to illustrate the
flexibility and potential of the power optimization method and
control optimization method.
[0247] FIG. 28 illustrates the main components of the circulation
system with the main BI-VAD components in place, with the right and
left side forcing functions from the native heart and the VAD. The
level of detail of system components shown in FIG. 28 is for
illustration purposes only, and several alternative models with
more or fewer details of system components can be drawn. One could
choose to obtain the dynamic response of the whole system,
including theoretical, numerical or experimental models for the
pulmonary system (lungs) and oxygenation dynamic systems. A
continuous model (differential equations expressing the dynamic
response at any position and time of the component) for all
components of the system would be of enormous complexity, though
still theoretically possible. Continuous models can also lead to
distributed parameter analyses. With current (2001) technology
those skilled in the art would likely (and straining computational
resources) choose to model: the L-VAD system from points M1 to M2;
the "right side" of the heart system from points M3 to M4 without
R-VAD; and from M3 to M5 with the R-VAD. These M1 to M5 points (or
other similar points that may be used for the optimization) are
used to separate suitable portions of the physical system in order
to develop a dynamic model of portions of the system rather than
the whole system shown in FIG. 28. Points such as M1 to M5 (or
other similar points) used in order to simplify the model, may
represent physical cross-sections of the flow passages, usually in
the main blood vessels. Those with ordinary skill in the art may
choose to develop the dynamic model using combinations of:
[0248] (a) discretized finite element method programs (FEM, for
example Szabo and Babuska, 1991; Bathe, 1995). For example, FEM
models may be used for the cardiac muscle, for structural
mechanical components, and other components;
[0249] (b) computational fluid dynamics models (CFD, for example
Anderson et al., 1984; Kiris et al., 1997). For example, CFD models
may be used for the blood and hydraulic fluid flows, and other
components of the system;
[0250] (c) analytic solutions for some dynamic elements. For
example, analytic models may be used for some of the fluid leakage
in narrow passages, using lubrication theory, and other
boundary-layer techniques. Samples of such methods for different
sub-components of the system are presented by (Nichols and
O'Rourke, 1998; Panton, 1984; White, 1991; Schlichting, 1979;
Hinze, 1987);
[0251] (d) specialized information for select parts of the
cardiovascular system. For example, select such models are
described by (Fung, 1984; Braunwald, 1984; Verdonck, 2000; Peskin
and McQueen, 1997);
[0252] (e) lumped parameter models for some of the dynamic
components. For example, some of the mechanical components may be
represented with techniques described by (Meirovich, 1975); and
[0253] (f) experimental data (which are usually the most reliable
models) can be used for, any aspect of the dynamic components of
the system.
[0254] Thus the overall dynamic model can be based on continuum
mechanics (or their variant of distributed parameter models), can
be discretized, can be based on lumped parameters, rely on
experimental data, or any combination thereof. Some of the
component models will be linear, others will be non-linear, and
some will be discrete or piecewise continuous (for example valve 70
is sometimes open, sometimes closed, and sometimes in the process
of opening or closing). The overall dynamic model is likely to be
complex, requiring significant computational resources. In
alternative embodiments useful information can also be derived from
simpler piecewise-continuous lumped-parameter non-linear dynamic
models for the main components. For example, such extremely
simplified models can consist of several masses, springs and
dampers for each of the main components shown in FIG. 28, so that
the overall dynamic model could be run on a conventional desktop
personal computer.
[0255] Once a suitable part of the physical prototype has been
defined with points such as M1 to M5 for development of the dynamic
model, the suitable inputs, outputs, boundary conditions, and other
system constraints must be carefully defined (this is referring to
the correct system definition, mentioned above). FIG. 29 is an
example of one possible dynamic-system representation of the main
components of the left side of the diseased native heart plus the
L-VAD. Alternative embodiments of the model may include portions of
the right side of the native heart and/or the R-VAD. If the sample
model for FIG. 29 extends between points M1 and M2 in FIG. 28, then
inlet and outlet boundary conditions would involve combinations of
blood pressures and velocities at M1 and M2 as functions of time.
In FIG. 29 the center diagram illustrates the dynamic model of the
physical system. The dynamic model can be developed with
experimental data or with equations or a combination thereof. There
are two basic physical components to the center block of FIG. 29.
One physical component is the diseased native heart and surrounding
tissue to the native heart; and the other physical component is the
VAD or BI-VAD or TAH system. The two physical components interact
and dynamically affect each other during normal operation as shown
by the dashed line R. The combination of the two main components of
the system, whether presented mathematically or with experimental
data, results in a dynamic representation of the physical system
from M1 to M2 that can be described by a form of equation (7). The
engineering definitions of the thermodynamic system (dynamic,
thermodynamic, fluid dynamic, mechanical, etc., as discussed
earlier) are crucial to the analysis and must be such that the
patient's tissue surrounding both the native heart and the VAD or
TAH are sufficiently removed from the components so that the
dynamic operation does not affect the boundary surface that
separates the dynamic system from the surrounding tissue. This is
an imaginary boundary surface typical of the boundary surfaces in
thermofluid dynamics texts that define the boundary of the
engineering system being analyzed. The input for the dynamic
representation of subsystems and components (shown on the left side
of FIG. 29) can come from different sources. Some of the input can
be experimental data, which is usually the preferred source of
data, but other input can come from other adequate mathematical
representations. Such inputs may come from the constitutive
relationships of cardiac muscle or other muscle, or of the
surrounding tissue, or the constitutive relationships for blood
flow, whether it is modeled as Newtonian or non-Newtonian,
incompressible or compressible. Several illustrations of these
models are published in the above references. The mathematical
representation of the physical system depicted in FIG. 29 from M1
to M2 can use constitutive relationships for the mechanical
components, or constitutive relationships for electromagnetic
components, or experimental data, or any combination thereof, or
experimental data and mathematical expressions of constitutive
relations, as shown on the left side of FIG. 29.
[0256] Other components of the physical system can also be
incorporated, as needed, depending upon how many or how few of the
system components are used in the dynamic model represented by the
final form of "process equations" (7) and (8). The level of detail
of the dynamic model will affect the complexity of the required
solution. It will also affect the fidelity and thus accuracy of the
results. In general, the higher the complexity and the fidelity the
more accurate the results, but at some point there is a limit;
i.e., a point of diminishing returns, where the increased
complexity does not justify the higher accuracy. A judgment must be
made on the fidelity of the dynamic model required for the
particular application of the invention. The final decision on this
issue will also depend on the sophistication of engineering tools
available, (CFD, FEM etc) and the level of accuracy required of the
results.
[0257] There are several alternative types of controllers suitable
for the application. In the simplest case the controllers may give
constant (battery) voltage, and vary the currents as a function of
time to the three coils of FIG. 18 (for example i.sub.1{t},
i.sub.2{t} and i.sub.3{t}). Other controllers may give constant
current but varying voltages. Other controllers may vary both the
voltage and the current. The latter is the most likely embodiment.
The third solution is likely to give the least power required than
the other two controllers, as the electrical resistance, inductance
and impedance of the coils, driving magnet 40, and surrounding
ferromagnetic material impose non-linear effects on the unsteady
flows of voltage and current, and the third type of controller
allows one to take full advantage of the "natural frequencies" of
the dynamic magnetic system in relation to the forcing function
required by the patient-VAD system.
[0258] The dynamic model developed above is used in the power
optimization method, an example of which is illustrated in FIG. 30.
The purpose of this optimization method is to minimize power
requirements and maximize battery life between recharges. This is
accomplished by identifying the minimum electrical power required
to the coils for each operating condition of the VAD. Since the
diseased native heart and VAD affect each other's dynamic
performance (broken line U in FIG. 29), the optimization process
must be repeated separately for each initial operating condition of
the unaided native heart. The inputs for the specific illustration
example are the initial condition of the diseased native heart,
comprised of the ECG trace and the ejected blood volumes and
pressures of the right and left side (atria and ventricles) as
functions of time over the period of the heart rate. In one
embodiment of the optimization process an intermediate output of
the power optimization method is X/Xmax {t} and F.sub.vad{t} of the
R-VAD and L-VAD. In an alternative embodiment of the optimization
process the output is the voltage and current fed to each coil as a
function of time as shown in FIG. 30. This optimization process
must be repeated for each initial condition of the diseased native
heart identified above.
[0259] One or more of the several potential optimal solutions V{t}
and i{t} are stored in controller 64. The choice of optimal
solution to insert in the controller is illustrated with an example
below.
[0260] The example power optimization method of FIG. 30 searches
for shapes of X/Xmax{t} that require the minimum power. In one
example embodiment of this optimization process, the ejection
volume of the native left ventricle is evaluated using MRI,
echocardiography, or other similar techniques such as correlation
with the movement of proximity sensors as described earlier. The
desired additional volume that must be provided by L-VAD 74 is
evaluated by methods illustrated in the earlier table of potential
sizes of VAD. This dictates the required travel X/Xmax of L-VAD 74
in FIG. 16. Next, an initial estimate for the trace of line 162 in
FIG. 17 (starting with an initial shape resembling that of line
162) is input into the power optimization method. This line shape
may also be modeled with Fourier series analysis, and the
amplitudes and phases in these Fourier series have a phase
difference from the ECG trace of the native heart in F.sub.nh{t}.
One measure of these phases is graphically reflected in the phase
difference .phi. from the phase of R in the QRS complex (phase
zero) to point 154 in the trace of X/Xmax in FIG. 17.
[0261] This concept is commonly referred to as "phase" in dynamic
systems. The pressures of the four chambers as functions of time
are required at least for the optimization sequence, and it may
also be required during the normal running of the device. However,
the pressures and the volumes can also be correlated by other means
in normal running of the device (for examples the ECG signal alone,
or the ECG signal plus volume traces, or ECG signal plus volume
plus pressure traces).
[0262] The forcing function of the native heart can be measured
(for example with measurements of ventricular and atrial pressures
and volumes, or their correlations) as described elsewhere in the
text. The forcing function of the VAD is an input to the
optimization process as described below. There are several
alternative combinations of specifying this forcing function as an
input to the optimization process. For example, one way is to
prescribe the displacement X/Xmax of driving magnet 40 as a
function of time, (FIGS. 16 and 17), evaluate the required force on
driving magnet 40 from the coils, and then evaluate the electrical
power required from the controller (voltages and/or currents to the
coils) to accomplish this motion. This is further elaborated below.
Then in an iterative process the displacement versus time (of
driving magnet 40) can be changed until the electrical power to the
coils is minimized while the displacement of driving magnet 40
provides corresponding displacements of driven magnet 54 that
result in acceptable ranges of volumetric blood throughput and
heart rate.
[0263] This initial estimate of the forced motion of X/Xmax {t}
results in changes in the pressure supplied by the combined
diseased left ventricle plus L-VAD 74. The result is that the
pressure and volume traces of the diseased heart shown in FIGS. 14
and 15 are modified, because the dynamic response of the native
heart system and L-VAD system affect each other. (In simple terms,
the motion of the one-way valve 70 from the aortic valve to the
aortic arch sucks additional blood per heart beat from that
accomplished by the diseased native ventricle alone, thus
increasing ejected blood volume per beat, so that the whole system
would tend to operate at lower heart rates).
[0264] The initial estimate of the shape of X/Xmax{t} results in a
required forcing function F.sub.vad{t} that must be provided by
driving magnet 40 to the hydraulic fluid and from there to the
blood, and this corresponds to the distribution of voltage and
current over time that the coils must provide to the driving magnet
40. The forcing function on driving magnet 40 is computed using the
dynamic model described above in equation (7). This forcing
function of the optimized design is compared with the 30-36 N
maximum force estimated in the discussion of FIGS. 1-4, above. The
electrical power (V{t} and i{t}) required to provide this force
(FIG. 17, line 180).is computed using dynamic models of the
transmission of power from the coils to the magnet, or measured
experimentally, or with a similar technique, reflecting equation
(8). The shape of Xmax {t} versus time is iteratively manipulated
until the electrical power required is minimized.
[0265] In alternative embodiments of the power optimization method
this minimization can be done numerically or experimentally, or by
neural networks to handle the volume of data and computations
required. The optimization of the transmission of electromagnetic
power can be done for at least three different cases, depending on
the type of controller 64: (a) the coils are supplied with constant
voltage and power changes are obtained with changes in the
electrical current; (b) The coils are supplied with constant
current and power changes are obtained with changes in the voltage;
and (c) the controller can vary both the current and voltage
applied to each coil as a function of time.
[0266] It is expected that for a given initial diseased heart
condition (for example, a heart rate of 100 beats per minute and
ejected volume from the unaided left ventricle 50 cc) the power
optimization process will result in several combinations of
modified heart rates and ejected volumes (from the combined left
ventricle and L-VAD) with slightly different power requirements.
For example, three potential L-VAD solutions, each with a different
shape of X/Xmax in FIGS. 16 and 17 to the above diseased-heart
condition, may be:
[0267] a. 80 beats per minute, 80 cc per beat, 7.0 Joules per beat
(560 Joules/minute);
[0268] b. 130 beats per minute, 60 cc per beat, 3.0 Joules per beat
(465 Joules/min);
[0269] c. 60 beats per minute, 90 cc per beat, 8.0 Joules per beat
(480 Joules/min).
[0270] In the last step of FIG. 30, for most practical applications
a cardiologist would choose to store in controller 64 either
solution (a) or solution (c) rather than the lower-energy solution
(b). These "optimal" solutions are obtained in an "external
optimization process" shown in FIG. 28 for a wide range of diseased
heart conditions and stored in controller 64. The combined output
of the diseased native heart and the new VAD/TAH is optimized both
for the individual diseased native heart of the patient and for the
power required to drive the artificial device. The output of the
power optimization method is the electrical power, and combinations
of voltage and current, that must be applied by the controller to
the coils.
[0271] In alternative embodiments this power optimization method
can be carried out mathematically (equations (7) and (8) for the
chosen system), or experimentally (with patient and VAD) in
clinical trials for groups of patients, or individuals patients. In
either case these optimization processes would benefit by the use
of neural networks.
[0272] The VAD installed in the patient must be able to adapt to
dynamic changes from one condition to the other, as the patient
with the VAD implanted in normal operative condition goes through
normal daily activities requiring changes in heart rates and
ejected blood volumes. The optimal design of the multi-input
multi-output dynamic system of the patient is illustrated in FIG.
31. In one alternative embodiment of the control optimization
method the physical dynamic system in FIG. 31 is the patient with
the VAD installed, or in other words the control optimization
method is experimental and is done clinically. In another
alternative embodiment the control optimization method is done with
the dynamic model of the physical prototype, reflected in
expressions of"process equations" (7) and (8).
[0273] Examples of the dynamic system shown in FIG. 31 are dynamic
models such as those shown in FIGS. 28 and 29, and incorporate the
results of the power optimization method of FIG. 30 (that defines
the steady-state, non-dynamically changing conditions of the
system). In control-system terminology this is a multivariate
control scheme (as opposed to more usual control schemes for
simpler linear mechanical systems). For purposes of this document,
"multivariate" means that the output state of the dynamic system is
characterized by several input variables and several output
variables, illustrated by the incoming and outgoing arrows on the
left and right side of FIG. 31.
[0274] Examples of these output-state variables are the ECG trace,
the blood volumes ejected from the ventricles, the flow rates
through points such as M1, M2, M3 and M5, the blood pressures or
hydraulic-fluid pressures at various points in the flow system,
other similar quantities, rates of change of these variables with
time, or combinations thereof. The state of these variables is
measured by various pressure, velocity, position, etc.,
transducers. Information about the output state variables is fed
back to the control node (x in circle) via the feedback transfer
matrices G(s). The purpose of the control optimization method
is:
[0275] (a) to find the optimum output variables for the control
scheme;
[0276] (b) to find the types and values of these feedback transfer
matrices G(s), which feed back signals to the control node; and
[0277] (c) to find the optimum input state variables for the
control scheme.
[0278] This optimization method is a multivariate input-output
control method with several input-output state variables. The input
state can be defined with variables such as beats per minute, the
phase .phi. of the distance from point Q to point 154 in FIG. 17
(in units of degrees or in units of time), the value of X/Xmax, the
shape of X/Xmax, several other similar quantities, or their rates
of change with time, and combinations thereof. In general, these
input-state control variables would be different from any similar
quantities that were computed in the power optimization process.
This does not preclude using the quantities from the power
optimization process, but this may lead to system instabilities
during transients.
[0279] Thus the inputs to the dynamic system are manipulated by
functions of the outputs of the dynamic system (which can be the
physical patient and VAD, or the dynamic model developed above) as
affected by the feedback transfer matrices G(s). The modified
inputs are fed into the dynamic system and affect its output state.
This optimization of the control method can be done analytically,
experimentally, or with various heuristic methods. Essentially
these control methods ensure that corrections to the dynamic system
to accommodate transient operations do not become unstable.
[0280] Simplified versions of this control sequence can be analyzed
with linear control theory. However it is more likely that
development of the control method will require well established
analytic and experimental techniques of non-linear, discrete or
continuous systems control, as several of the elements of the
dynamic system (e.g. in FIG. 29) are non-linear.
[0281] Suitable analytic optimal control techniques have been
published in the open literature (Brogan, 1990; Glad and Ljung,
2000; Fradkov et al., 1999; Schroder, 2000). However, it is
envisioned that in alternative embodiments neural networks,
adaptive control techniques, and observer-based methodologies will
be suitable alternative embodiments of the new control optimization
method. The final step in the control optimization sequence is to
store the optimized feedback transfer matrices G(s), and associated
control scheme, into controller 64.
[0282] An additional way to illustrate the flow of information flow
in the new device during normal operation is shown in the flow
charts of FIGS. 32, 33 and 34, all of which include points M1, M2,
M3 and M5 that are also in FIGS. 28 and 29. FIG. 32 shows the
application to an L-VAD. The figure shows at the top the native
heart as right and left sides, atria and ventricles, and in the
center there is an illustration of the ECG signal, which feeds
information to the controller. In this case, shown in FIG. 32, the
information would be the ECG trace, or heart rate and phase
entering the L-VAD controller. (Alternative embodiments may include
measures of volumes and pressures in atria and/or ventricles). The
left ventricle provides some output that goes into the control
junction, shown by an X in a circle, which is also fed into the
controller. The controller, using information provided by the power
optimization method and control optimization method that were shown
and discussed with reference to FIGS. 30 and 31, provides electric
power, V{t} and i{t} to the coils, that dictates the movement of
driving magnet 40. Driving magnet 40 in turn dictates the movement
of driven magnet 44, and that in turn dictates the movement of
valve-seat magnet 54, which affects the output of the left
ventricle (point M2). A similar arrangement could be drawn for the
R-VAD, but it is exactly symmetrical to the one shown here so this
is not drawn or described further.
[0283] The flow chart of FIG. 33 illustrates the application of the
new optimization process to a bi-ventricular assist device (BI-VAD,
described above). The flow chart is split into two parts, for the
L-VAD and the R-VAD, right and left respectively. The components
themselves and the logic flow paths are similar to those shown in
FIG. 32, and thus are not discussed further herein.
[0284] FIG. 34 shows the application of the new optimization
process to the total artificial heart (TAH) embodiment, in which
the left and right ventricles are removed, but signals are still
received from the sinoatrial node. These signals are provided to
the L-TAH and R-TAH controllers 64. In response, the controllers
drive magnets 40, 44 and 54, and finally these magnets provide the
overall volumetric throughput for the cardiac system, corresponding
to points M2 and M5 in FIGS. 30 and 31, as previously
discussed.
[0285] These mathematical and engineering techniques will be
augmented by clinical trials on groups of patients, and
standard-sized or unique VAD devices sized for individual patients
devices may be optimized to the individual patients. As the
condition of the patient changes with time, the control variables
and control scheme stored on controller 64 will need to be updated
to the new condition of the patient. The power and control
optimization sequences are identical to the sequences described
above. The new data for controller 64 can be transmitted to the
controller inside the patient's body using established infra red
data transmission techniques, or other similar techniques.
[0286] Several alternative embodiments to the described systems and
methods are also conceived. For example, one alternative embodiment
entails the use of neural networks, or comparable technology, to
optimize the displacements shown in FIG. 28 (X/Xmax and their
shape), with the ejected blood volume. For example, with reference
to FIG. 29, dynamic measurement of volume ejected and phase may be
eliminated, because volume and phase can be correlated (with neural
networks) with the motion of two or more proximity sensors, as
shown in FIG. 27.
[0287] Alternative embodiment of the above-described methods are
conceived in which the optimization processes for FIGS. 30 and 31
is not done mathematically, but it largely depends on clinical
trials with extensive use of neural networks to expedite the
computation process.
[0288] Another enhancement of the new system is that the controller
can detect the presence of certain arrhythmias, such as ventricular
tachycardia, for example. In this event, the electromagnetic pump
coils would be de-energized and no current would be supplied to
such coils, as it would be undesirable for the VAD to be activated.
As a "fail-safe", if the VAD was in fact de-energized, in such a
case, the one-way valve 70 inside of valve-seat magnet 54 would
respond to the pressure gradient of the blood flowing past it, and
would open and close via the forces applied to it by the flowing
blood.
[0289] In view of the foregoing, it will be seen that the several
objects of the invention are achieved and other advantages are
attained. Although the foregoing includes a description of the best
mode contemplated for carrying out he invention, various
modifications are conceivable. As various modifications could be
made in the constructions and methods herein described and
illustrated without departing from the scope of the invention, it
is intended that all matter contained in the foregoing description
or shown in the accompanying drawings shall be interpreted as
illustrative rather than limiting.
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