U.S. patent number 3,698,381 [Application Number 05/102,778] was granted by the patent office on 1972-10-17 for monitoring system for physiological support systems.
This patent grant is currently assigned to Avco Corporation. Invention is credited to Armando Federico, Henry R. Guarino.
United States Patent |
3,698,381 |
Federico , et al. |
October 17, 1972 |
MONITORING SYSTEM FOR PHYSIOLOGICAL SUPPORT SYSTEMS
Abstract
A system for automatically monitoring and controlling the fluid
load in physiological support systems. The system samples gas
pressure at fixed periods and adds fluid or vents fluid from the
system to prevent malfunction and the catastrophic results.
Inventors: |
Federico; Armando (Medford,
MA), Guarino; Henry R. (Revere, MA) |
Assignee: |
Avco Corporation (Cincinnati,
OH)
|
Family
ID: |
22291632 |
Appl.
No.: |
05/102,778 |
Filed: |
December 30, 1970 |
Current U.S.
Class: |
600/17; 600/18;
604/914 |
Current CPC
Class: |
A61M
60/50 (20210101); A61M 60/268 (20210101); A61M
2205/33 (20130101); A61M 2205/3303 (20130101); A61M
60/40 (20210101); A61M 60/135 (20210101); A61M
60/274 (20210101) |
Current International
Class: |
A61M
1/10 (20060101); A61b 019/00 () |
Field of
Search: |
;128/1R,DIG.3
;3/1,DIG.2 |
References Cited
[Referenced By]
U.S. Patent Documents
Primary Examiner: Trapp; Lawrence W.
Claims
What is claimed is:
1. In a circulatory assist system which sequentially actuates a
blood pump by controlling fluid load pressure to the pump, the
combination comprising:
a. sensing means for sensing pump fluid load pressure;
b. first means responsive to said sensing means for increasing said
pressure when it falls below a first predetermined level; and
c. second means responsive to said sensing means for venting said
pressure to ambient pressure when the load pressure exceeds a
second predetermined level.
2. The combination as defined in claim 1 and further including
third means responsive to said sensing means for venting said
pressure to ambient pressure when said pressure falls below a third
predetermined level.
3. The combination as defined in claim 2 and further including
first timing means responsive to said first means for measuring a
first predetermined time interval and further including fourth
means responsive to said first timing means and said sensing means
for venting said pressure to ambient pressure when said pressure
increasing to a fourth predetermined level does not occur within a
first predetermined interval.
4. The combination as defined in claim 3 and further including a
second timing means responsive to said sensing means for measuring
a second predetermined time interval and further including fifth
means responsive to said second timing means and said means for
venting said pressure to ambient pressure when said pressure falls
from above a fourth predetermined level to below said first
predetermined level within a second predetermined interval.
5. The combination as defined in claim 4 wherein said fourth
predetermined level is greater than said first predetermined level
and is less than said second predetermined level.
6. The combination as defined in claim 2 wherein said third
predetermined level is less than said first predetermined level
which is less than said second predetermined level.
7. The combination as defined in claim 4 wherein the first
predetermined interval is longer than the second predetermined
interval.
8. The combination as defined in claim 4 and including further
means responsive to said sensing means for changing pressure when
it deviates from a normal level within a predetermined pressure
range lying between said first predetermined level and a fifth
predetermined level.
9. The combination as defined in claim 8 wherein said last
mentioned pressure change in response to load pressure exceeding
the normal level within the predetermined pressure range is
accomplished by relatively slowly venting to ambient as compared to
the rate of venting when load pressure exceeds the second
predetermined level.
10. The combination as defined in claim 8 wherein said fifth
predetermined level is greater than said first predetermined level
and is less than said second predetermined level.
11. The combination as defined in claim 2 wherein said first means
includes a slow fill valve whose rate is not substantially in
excess of in vivo blood to absorb the fluid utilized.
12. A circulatory assist system comprising:
a. Fluid driven pump means for pumping blood;
b. means for sensing pump fluid load pressure;
c. means for venting said load pressure to ambient;
d. means for increasing said load pressure; and
e. a plurality of means responsive to said sensing means including,
first means for controlling said venting means within a
predetermined load pressure range, second means for controlling
said pressure increase means within said same predetermined load
pressure range, and third means for controlling said pressure
venting means above and below said predetermined load pressure
range.
13. A system as defined in claim 12 wherein, the venting means
includes a relatively slow vent and a relatively fast vent, the
latter serving to vent above and below the predetermined load
pressure range.
14. A system as defined in claim 13 further including fourth means
coupling the first and second controlling means to said low
pressure venting and to said increasing means, thereby causing
pressure to vent when the pressure falls below said predetermined
range within a first predetermined interval.
15. A system as defined in claim 14 and further including fifth
means coupling said first and second controlling means to said low
pressure venting for said increasing means, thereby causing
pressure to vent when the time required to increase pressure to
said predetermined range exceeds a second predetermined
interval.
16. A system as defined in claim 15 wherein the first predetermined
interval is longer than the second.
Description
This invention relates to physiological support systems which
require a membrane separation between blood and a fluid. These
systems fall into two categories. The first type is a mechanical
device to support cardiac action such as left ventricular assist
device. In this type of support system pressure is exerted upon and
transmitted across a membrane in contact with the blood. Blood is
displaced with the creation of controlled pressure changes which
tend to reduce cardiac effort and pump blood systemically. The
membrane which separates the driving gas or liquid from the blood
in these devices is subjected to repeated bending stresses and
requires superior mechanical properties. The second general
category of devices are diffusion systems. Examples of these are
the membrane oxygenator type of artificial lung and the
hemodialysis type of artificial hiding. In these types of devices
the membrane serves to separate the blood from the diffusion gas or
liquid and is permeable to the substances which are exchanged from
fluid to flood or vice versa. The separating membrane must be
generally quite thin in relation to its surface area. One possible
hazard of both categories of devices is the development of gas
leaks and the embolization of fluid into the bloodstream. The
physiologic consequences of such an event depend on the amount and
type of fluid embolized. The body's tolerance for most foreign
fluids in bulk amount is not large. Therefore, it may now be seen
that major leaks in most if not all cases would be a catastrophic
event and that a patient must be protected against such an
event.
It is one purpose of the present invention to provide an improved
heart assist apparatus which augments blood flow during diastole
and decreases myocardial effort during systole.
Heretofore, mechanical assistance to the failing heart has been
attempted by veno-arterial pumping, arterio-arterio pumping, and a
variety of counterpulsation techniques including intra-aortic
balloon pumping. In the counterpulsation, it is necessary for the
pump to be synchronized with the patient's heart. Furthermore, most
if not all of these systems include a fluid powered pump. The fluid
is usually gas and exerts pressure on a flexible member or
membrane, which in turn exerts a force on the blood. Thus, the
fluid or gas load pump is separated from the blood and great care
is taken to insure against leakage of gas into the blood or leakage
of blood into the gas load system.
It is another object of the present invention to provide automatic
load fluid or gas surveillance for such systems.
It is another object to provide a monitoring system for such heart
pumps that will be fail-safe, reducing danger to the patient.
It is a further object to provide a monitoring system for such
heart pumps that will cause the pump to be fail-safe, in the event
of pressure extremes in the fluid load pressure and/or excessive
leakage between the blood and the load fluid.
One type of heart pump, the intra-arterial balloon pump and more
particularly the intra-aortic balloon pump, consists of a balloon
attached to a catheter, which is introduced into the patient's
femoral artery and positioned in the descending aorta to provide
benign cardiac assistance to the patient. A constant amount of load
gas is required in the system at all times to insure proper
inflation and deflation of the balloon which is controlled by a
volumeter that increases and decreases the pressure on the load
gas. However, due to leaks, diffusion, and accidental malfunction
of the operating parts, gas may leak from the system. Excessive
amounts of gas leaking into the patient's blood system must be
stopped to protect the patient. Also, excessive leaks from the gas
system outside the patient's body, which indicate a substantial
failure in the system must be stopped. On the other hand, minor
leaks due to osmosis through the balloon or the catheter, or minor
leaks from the system outside the patient's body, or slight
increases in pressure due to pinched catheter line can be remedied
by filling or venting the load gas to maintain proper operating
pressure.
It is another object of the present invention to provide a gas
surveillance system for such an intra-aortic balloon pump, which
meets at least some of these safety requirements.
It is another object that the surveillance system be fail-safe so
that the balloon pump in the patient's body imposes the least
restriction to blood flow upon failure.
In accordance with the present invention, the gas pressure load in
an intra-arterial or intra-aortic pump is measured during quiescent
periods between pump impulses and the load system is filled or
vented slowly if pressure is within a safe region to maintain ideal
pressure load. If pressure is outside the safe region, the system
is rapidly vented to collapse the balloon so that it will provide
minimal resistance to blood flow. The surveillance system also
vents to collapse the balloon if filling to maintain the desired
pressure occurs too frequently or if it takes too long to fill to
maintain normal operating pressure. Either of these indicates there
is excessive leakage somewhere in the load system.
The novel features that are considered characteristic of the
present invention are set forth in the appended claims. The
invention itself, however, both as to its organization and method
of operation, together with other objects and advantages, will be
best understood from the following description of a specific
embodiment of the invention, taken in conjunction with the
accompanying drawings, in which:
FIG. 1 is a block diagram and schematic showing the pneumatic gas
load system, the intra-aortic pump and the electronic control and
surveillance system which responds to, for example,
electrocardiograph equipment monitoring the patient;
FIG. 2 is a block diagram showing the circuits in the rapid vent
and refill logic of the surveillance system; and
FIG. 3 shows waveforms in the surveillance system, including a
waveform representative of system load pressure and various levels
identified with safe and unsafe operation.
Turning first to FIG. 1, there is shown a block diagram and
schematic of a pneumatic system and an electronic control and
surveillance system, which function together to control an
intra-aortic balloon 1, attached to a catheter 2, which is
introduced into the patient's femoral artery and positioned in the
descending aorta to provide a benign cardiac assist to the patient.
The pneumatic system consists principally of two separate pneumatic
systems separated by the membrane 3 in isolating piston 4. It is
through the isolating piston, and most particularly the membrane 3
in the isolating piston that pressure pulses are delivered from the
initiating pneumatic system 6 to the pump load pneumatic system 7.
The gases in the systems 6 and 7 may be different because the
requirements of the systems are different. For example, the pump
load system which extends into the body of the patient through the
intra-aortic balloon 1 preferably contains helium which dissolves
easily in the blood without injurious effect on the patient. Other
gases which are not substantially injurious are carbon dioxide and
oxygen and while these gases are heavier they may also be used
because there is a tolerance by the patient to dissolve these gases
in the blood. Nitrogen, on the other hand, could cause damage or
injury to the patient, because of the low toleration for nitrogen
in the blood. These considerations must be made because of the
possibility of gas leaking from the catheter or the balloon into
the patient's blood. Such considerations are not necessary with
regard to the gas load in the initiating pneumatic system 6 because
there is little possibility of it entering the patient's blood
system.
Clearly, the patient's toleration for the gas, which loads the pump
pneumatic system 7, will be one of the factors to determine when
gas leakage from the system could be harmful to the patient and so
constitute a malfunction and require shutdown or venting. The
purpose of the electronic control and surveillance system, shown in
FIG. 1, is twofold. First, it provides electrical trigger to the
initiating pneumatic system 6 that drives the isolating piston 4;
and second, in the pump load system 7 it responds to the pressure
of the load gas in the intra-aortic balloon and by providing a
signal to the pneumatics which maintains that pressure within
predetermined limits, refills the system where a leakage not
excessively harmful to the patient has occurred or vents the system
shutting it down where the leakage has been excessive or where
pressure is beyond prescribed pressure limits, indicating a serious
malfunction in the system. The system shuts down by venting so as
to collapse the intra-aortic balloon. The balloon is collapsed so
that it is least obstructive to blood flow in the aorta.
The intra-aortic balloon pump must be synchronized with the
patient's heartbeat. More particularly, the balloon must be pulsed
or inflated with pressure during the diastolic phase of the heart
cycle. During this phase, as already described, the left ventricle
of the heart which feeds the aorta is in dilation, the aortic valve
is closed and the aorta is normally distended. At this point, the
balloon is inflated and so it aids the aorta to force the blood
through the capillaries and other vessels of the patient's body. A
representation of the pressure pulses to the balloon is shown by
the waveform in FIG. 3 designated System Pressure. The duration I
of a pulse in this waveform represents the pumping interval and the
point S between pulses indicates a sample point when the pressure
to the balloon is sampled and examined by the surveillance system,
to determine whether it is within prescribed limits or a dangerous
malfunction threatens. The point S in the pumping cycle represents
the quiescent or zero pressure, which loads the pump between
pumping intervals.
The pumping interval is approximately two-thirds of the whole cycle
when the patient's pulse rate is 77 beats per second. This interval
may be reduced to about one-half the pump full cycle when the
patient's pulse rate is between 100 and 120 beats per second. An
electrical system which responds to the output of an
electrocardiograph (EKG) unit monitoring the patient's heartbeat
provides an electrical drive signal for controlling the pneumatic
system 6 that initiates a pressure pulse to the balloon and also
triggers a sampling pulse for sampling balloon pressure. This
system is designated the heart-pump sync system 10 and produces a
drive pulse which is positioned in the heart diastole phase,
depending on the patient's pulse rate. A suitable heart-pump sync
system of this sort is described in U.S. Pat. No. 3,452,739, which
is assigned to the same assignee as the present invention. The
signal output of the EKG is represented by the waveform denoted EKG
input, shown in FIG. 3, and the drive signal output from the
heart-pump sync system 10 is represented by the waveform designated
Drive Signal in FIG. 3.
The drive signal is amplified by amplifier 11 and energizes the
solenoid valve 12 in the pneumatic system 6. This valve connects
the isolating piston 4 to a pressure source 13 or a vacuum source
14 and so distends the diaphragm 3 in the isolating piston to
deliver pressure pulses to the pump load in the pneumatic system 7.
The pressure pulses are transmitted from the piston 4 to the
intra-aortic balloon pump 1 via the catheter 2.
Pressure to the pump is monitored by a pressure transducer 15 in
contact with the catheter. The electrical signal from the
transducer is amplified by amplifier 16 and fed to the electronic
surveillance system, denoted generally by the numeral 17. This
signal, the system pressure signal, is shown by the waveform in
FIG. 3, designated System Pressure. In the surveillance system the
system pressure signal is sampled during the interval of Sampling
Gate pulses, also shown in FIG. 3. These pulses are generated by a
monostable multi-vibrator 18, in response to the Drive Signal
pulses from the heart-pump sync system 10.
In operation, the pneumatic system 7 is initially filled with load
gas through valve 21 from a tank of helium gas 22, through a
pressure regulator 23. The load gas reservoir 24 in the isolating
piston 4 is filled to a maximum gauge pressure of 15 mm, as set by
the gas load regulator 23. This amount of initial load gas in the
reservoir is variable depending on the size of the balloon used,
type of load gas, and the length of the catheter. When the balloon
is not pumping, it is deflated and the pressure in the catheter is
established at safe operating limits for the system, which are by
way of example between +7 and -5 mm of pressure for helium. When
the system is pumping, the surveillance system 17 samples the
pressure between pump pulses and commands corrective action. If the
sampling is within the limits +7 to -5 mm no action is taken.
However, if the sample is not within these limits, corrective
procedures are introduced to insure proper and safe operation of
cardiac device implanted in the patient. If an unsafe over pressure
condition exists, for example; at any one sample the pressure is
+20 mm or greater when helium is used, the surveillance system 17
actuates rapid vent valve 25, which quickly vents the pump
pneumatic system to atmosphere thereby relieving any gas stored in
the isolating piston. This also shuts off the initiating pneumatic
system 6 by means which are not shown. If the pressure is in the +7
or +20 mm range, a slow vent valve 26 is opened, releasing the
excess gas to atmosphere with a controlled release rate until a
pressure of below +7 mm is achieved. Restoration to a safe
operating region is normally accomplished within a few heart
cycles. If the pressure drops to a low pressure, a typical range
for example being between -5 and -20 mm, refill load valve 27 is
actuated to replenish the gas load in pneumatic system 7. However,
before refilling, the surveillance system 17 asks if the time from
the last high normal pressure of +5 mm to which the system is
refilled has occurred within the last minute. If it is greater than
a minute, refill load valve 27 is opened to replenish the gas load
to the high normal pressure of +5 mm. However, if the refill valve
is opened continuously for more than 10 seconds, it is assumed
there is a leak present in the system and so the refill valve 27 is
shut and the rapid vent valve 27 is opened, shutting down the
system just as when an over pressure exceeding +20 mm occurs. On
the other hand, if the refill load valve 27 is open for less than
10 seconds, and the system is restored to the high normal operating
pressure of 5 mm. The system is allowed to continue pumping. When
the system pressure drops -20 mm or less, rapid vent 25 is also
opened as this indicates that a failure has occurred and the
patient is in danger.
In the surveillance system 17, signal compare circuits 31 to 35,
which may be commercially available voltage comparators, compare
the voltage of the pressure signal from amplifier 16 with preset
voltages representing the pressure limits designated and when the
pressure signal exceeds, or is less than, the designated limit
called for by the compare circuit, the circuit produces an output
signal. More particularly, circuit 31 produces an output signal
when the pressure signal exceeds a reference representing +7 mm
pressure. Circuit 32 produces an output when the pressure signal
exceeds a reference representing +20 mm of pressure. Circuit 33
produces an output when a reference signal representing -20 mm
exceeds the pressure signal. Circuit 34 produces an output when the
pressure signal exceeds the reference signal representing 5mm of
pressure and circuit 35 produces an output when a reference signal
representing -5 mm exceeds the pressure signal. The outputs from
circuits 31 to 35 are gated by gates 36 to 40 and the gates are
controlled by Sampling Gate pulses shown in FIG. 3 generated by
multivibrator 18. Thus, the outputs of the gates are binary signals
representing the occurrence of the events indicated during the
sample gate intervals, which fall between the system pressure
pulses as shown by waveforms in FIG. 3.
The output of gate 36 is a negative voltage pulse representing the
occurrence of the condition determined by circuit 31. The negative
pulse occurs if the condition occurs, such as, for example, a
pressure exceeding a reference pressure of +7 mm. The negative
pulse initiates a slow vent control signal from pulse shaper 41
amplified by slow vent amplifier circuit 48, which controls the
vent valve 26. This slow vent control signal is terminated when the
output of gate 36 is a positive level when the system is below +7
mm and is determined by AND circuit 42. Thus, the slow vent valve
26 vents the system continually following initiation and until the
system pressure is decreased below +7 mm and a sampling gate pulse
disappears.
The outputs of gates 37 and 38 represent the occurrence of an
extreme condition, for example, a pressure over or under + 20 mm
requiring rapid venting of the system and so these gates trigger
flipflop circuits 44 and 45, which feed through OR gate 46 to
amplifier 47 that energizes rapid vent valve 25. This shuts down
the system as danger points have been exceeded. Thereafter, when
the defects are corrected, flipflop 44 and 45 are reset and ready
again to initiate rapid vent in case pressure exceeds the extreme
limits.
The outputs from gates 39 and 40 trigger rapid vent and refill
logic circuits 50. These circuits feed signals through OR circuit
46 to the rapid vent amplifier 47 to cause rapid venting of the
system in accordance with the logic already described and also
perform the logic to determine whether the system will be refilled.
If refill is called for, then a signal from circuits 50 is
amplified by refill amplifier 51 which energizes refill load valve
27.
The logic circuits in 50 are shown in FIG. 2, which is a simple
block diagram of logic elements. This consists of a 1 minute timer
52 which responds to the output of gate 39 producing a pulse of 1
minute interval, when the pulse output from gate 39 is positive
indicating a high normal pressure of +5 mm. Similarly, the output
from gate 40 initiates a pulse at a low normal pressure of -5 mm.
The logic in circuits 54, which includes an AND circuit 55 and an
AND circuit 56 questions first whether the lower pressure limit -5
mm, which requires refilling, has been reached and then asks
whether one minute has elapsed since the pressure was last above +5
mm. If the answer to both of these is no, it means the system is
losing pressure too rapidly and so there is a serious leak and the
system must be shut down. The output of AND circuit 55 is fed
through the OR circuit 46 to energize the rapid vent valve 25 that
initiates shutdown of the system. On the other hand, if the answer
is yes, meaning that refill is required and the pressure has not
been above +5 mm during the last minute, then the output of AND
circuit 56 generates a fill signal to the flipflop 59. The output
of the flipflop 59 and a 10 second delay pulse from timer 53 are
combined by AND circuit 57. The logic here asks the further
question in response to a yes answer from circuit 54 "Is the system
still filling after 10 seconds?" If the answer to this is yes, then
the indication is that the system has not been restored to an
operating condition and it must be shut down. Thus a yes answer in
the output of AND circuit 57 commands the system to vent and shut
down because refilling is taking too long. However, if the pressure
reaches a high normal of +5 mm within 10 seconds, the filling is
stopped. This logic is further implemented by double input flipflop
circuit 59. One stage of the flipflop controlled by AND circuit 56
feeds through OR circuit 46 to control rapid vent valve 25 and the
other feeds through refill amplifier 51 to control refill valve
27.
It is envisioned that the timers mentioned above may be eliminated
along with the high normal comparator (+5 mm) circuit if it is
desired to use an analog refill system and/or it is desired to
eliminate the 1 minute and 10 seconds timing functions along with
the high normal pressure functions as described hereinabove. Such a
circuit is activated whenever the pressure drops to the low normal
level of -5 mm pressure. This will automatically activate the
refill load valve. The rate of refilling is made at least
substantially equal to the rate at which the human body may safely
absorb the load gas utilized. If this rate of refill is inadequate
to maintain the normal operating pressure and the system drops
below the safe limit of -20 mm pressure, the system will rapidly
vent and shut down as previously described. Further, if the refill
rate is sufficient to bring the system to within the safe operation
range, the patient will not be endangered because the rate of
adding the gas can be safely absorbed by the human body. If the
fill rate increases the pressure to the upper limit of the system,
a feedback loop or the like shuts down the fill function at the
desired high normal pressure of, for example, +5 mm.
The embodiment of the invention described herein both as to general
and specific functions and general and specific structures
illustrates the best known use of the invention. It will be
apparent to those skilled in the art that the invention has use
also in control and surveillance of pressure in other types of
circulatory assist pumps, where the pump is located inside or
outside of the patient.
* * * * *