U.S. patent application number 13/946337 was filed with the patent office on 2015-01-22 for system and method for monitoring cardiac output, flow balance, and performance parameters.
The applicant listed for this patent is Mohammad Khair. Invention is credited to Mohammad Khair.
Application Number | 20150025328 13/946337 |
Document ID | / |
Family ID | 52277615 |
Filed Date | 2015-01-22 |
United States Patent
Application |
20150025328 |
Kind Code |
A1 |
Khair; Mohammad |
January 22, 2015 |
SYSTEM AND METHOD FOR MONITORING CARDIAC OUTPUT, FLOW BALANCE, AND
PERFORMANCE PARAMETERS
Abstract
A system for measuring of cardiac output and cardiac performance
parameters based on a cardiac blood flow balance parameter between
a right chamber of the heart and a left chamber of the heart,
includes a sensor device for measuring one of blood pressure and
blood flow rate and blood constituent concentration of a patient so
as to generate an arterial pulse signal. A processing unit is
responsive to the arterial pulse signal for generating a full
arterial pulse signal, an arterio-venous pulse signal, and a
balance parameter. A computational device is responsive to the
balance parameter for further generating cardiac output and a set
of cardiac performance parameters. A display station device is
responsive to the set of physiological parameters from the
computational device for displaying meaningful information.
Inventors: |
Khair; Mohammad; (Irvine,
CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Khair; Mohammad |
Irvine |
CA |
US |
|
|
Family ID: |
52277615 |
Appl. No.: |
13/946337 |
Filed: |
July 19, 2013 |
Current U.S.
Class: |
600/301 ;
600/501; 600/526 |
Current CPC
Class: |
A61M 1/1086 20130101;
A61B 5/029 20130101; A61B 5/0295 20130101; A61B 5/7225 20130101;
A61M 2230/005 20130101; A61M 1/12 20130101; A61M 2205/3303
20130101; A61B 5/14546 20130101; A61B 5/7239 20130101; A61B 5/0261
20130101; A61M 1/10 20130101; A61N 1/3987 20130101; A61B 5/7475
20130101; A61B 5/02055 20130101; A61M 1/14 20130101; A61M 2230/42
20130101; A61B 5/742 20130101; A61M 16/01 20130101; A61B 5/7257
20130101; A61B 5/14532 20130101; A61M 5/1723 20130101; A61B 5/14551
20130101; A61B 5/026 20130101; A61B 5/0402 20130101; A61B 5/726
20130101; A61N 1/36585 20130101; A61B 5/14552 20130101; A61B 5/0215
20130101; A61M 1/1006 20140204; A61B 5/14542 20130101; A61M 2205/33
20130101; A61M 2205/3334 20130101; A61B 5/02028 20130101; A61B
5/0205 20130101; A61M 1/122 20140204; A61B 5/021 20130101; A61B
5/02108 20130101; A61B 5/02433 20130101 |
Class at
Publication: |
600/301 ;
600/501; 600/526 |
International
Class: |
A61B 5/00 20060101
A61B005/00; A61B 5/02 20060101 A61B005/02; A61B 5/021 20060101
A61B005/021; A61B 5/0402 20060101 A61B005/0402; A61B 5/029 20060101
A61B005/029 |
Claims
1. A method for determining cardiac performance parameters so as to
assess a patient's health condition, comprising the processing
steps of: measuring of one of an aortic pressure signal and a
pulmonary arterial pressure signal; measuring of an
electrocardiogram or electrogram waveform signal; computing of a
first derivative of the measured aortic or pulmonary arterial
pressure signal; determine a heart valve opening event using the
maximum peak of the first derivative of the measured aortic or
pulmonary arterial pressure signal; determine a heart valve closing
event using the minimum valley of the first derivative of the
measured aortic or pulmonary arterial pressure signal; optionally
determine an end systolic pressure event and value using the
maximum of the measured one of aortic pressure, pulmonary arterial
pressure, and ventricular pressure signal; calculating a full
arterial pulse waveform signal from the said measured aortic or
pulmonary arterial pressure signal, defined as the aortic or
pulmonary artery pressure signal without the effect of atrial
diastolic blood flow demand; calculating an arterio-venous pulse
waveform signal by the subtraction of the said measured aortic or
pulmonary arterial pressure signal from the said full arterial
pulse waveform signal; determining a pre-diastolic end-systolic
pressure (PDPES) event, defined by the ventricular pressure post
the isovolemic relaxation phase and prior to diastolic expansion,
which is substantially coincident in time with the end of the
arterio-venous pulse; optionally calculating a Left-Right Balance
or Right-Left Balance parameter based upon the full arterial pulse
and the arterio-venous pulse waveform signals; determining an
end-diastolic pressure (PED) event, defined by the ventricular
pressure post diastolic expansion and pre-isovolemic contraction
phase, which is substantially coincident in time with the R-wave
event in the electrocardiogram or electrogram waveform signal;
calculating a first timing differential between end-diastolic
pressure (PED) event and the valve opening event; estimate the end
diastolic pressure (PED) value using said first timing differential
in extrapolating backwards a first line segment of the measured
aortic or pulmonary arterial pressure signal between the valve
opening event and short time thereafter; calculating a second
timing differential between the said pre-diastolic end-systolic
pressure (PDPES) event and the valve closing event; estimate the
said pre-diastolic end-systolic pressure (PDPES) value using said
second timing differential in extrapolating forward a second line
segment of the measured aortic or pulmonary arterial pressure
signal between the valve closing event and short time
therebefore;
2. A method for determining cardiac performance parameters as
claimed in claim 1, further including the step of using the end
diastolic pressure (PED) and the pre-diastolic end-systolic
pressure (PDPES) values to assess cardiac ventricular end diastolic
volume and end systolic volume.
3. A method for determining cardiac performance parameters as
claimed in claim 2, further including the step of using the end
diastolic pressure and the pre-diastolic end-systolic pressure
(PDPES) values to calculate ejection fraction and using cardiac
ventricular end diastolic volume and end systolic volume to
calculate cardiac output.
4. A method for determining cardiac performance parameters as
claimed in claim 3, wherein the ejection fraction EF is equal to
1-PDPES/PED, where PDPES is the pre-diastolic end-systolic pressure
value and PED is the end-diastolic pressure value.
5. A method for determining cardiac performance parameters as
claimed in claim 4, further including the step of determining
cardiac ventricular end diastolic volume and end systolic volume
based upon measuring the ventricular volume at end-systolic and
end-diastolic events for a calibration.
6. A method for determining cardiac performance parameters as
claimed in claim 5, further including the step of calculating the
volume to static pressure conversion as a linear scale ratio R
defined by (VED-VES)/(PED-PDPES), where PED is the end diastolic
pressure value, PDPES is the pre-diastolic end-systolic pressure
value, VED is the cardiac ventricular end diastolic volume and VES
is the cardiac ventricular end systolic volume.
7. A method for determining cardiac performance parameters as
claimed in claim 6, further including the steps of continuously
estimating the end diastolic volume VEDest to be equal to R*PED and
continuously estimating the end systolic volume VESest to be equal
to R*PDPES.
8. A method for determining cardiac performance parameters as
claimed in claim 7, further including the step of continuously
estimating stroke volume SV to be equal to the difference between
the estimated end diastolic volume and the estimated end systolic
volume VEDest-VESest.
9. A method for determining cardiac performance parameters as
claimed in claim 8, further including the step of continuously
estimating cardiac output CO to be equal to the stroke volume
SV*HR, where HR is the heart rate.
10. A method for determining cardiac performance parameters as
claimed in claim 9, further including the step of calculating
cardiac output index CI to be equal to the cardiac output CO
divided by BSA, where BSA is the body surface area.
11. A method for determining cardiac performance parameters as
claimed in claim 1, further including the step of calculating mean
arterial pressure MAP.
12. A method for determining cardiac performance parameters as
claimed in claim 11, further including the step of calculating
systemic vascular resistance SVR to be equal to the difference
between the mean arterial pressure MAP and a central venous
pressure CVP, which is then divided by the cardiac output CO.
13. A method for determining cardiac performance parameters as
claimed in claim 8, further including the step of calculating
stroke work SW to be equal to the stroke volume SV times the mean
arterial pressure MAP.
14. A method for determining cardiac performance parameters as
claimed in claim 3, further including the step of calculating
contractility by using the derivative of the ejection fraction
EF.
15. A method for determining cardiac performance parameters as
claimed in claim 6, further including the step of using the linear
scale ratio R computed for one ventricular chamber's volume to
estimate the other ventricular chamber's volume and its cardiac
performance parameters.
16. A method for determining cardiac performance parameters as
claimed in claim 1, further comprising the steps of providing a
processor unit and configuring the processor unit to perform all
said processing steps.
17. A method for determining cardiac performance parameters as
claimed in claim 16, wherein said processor unit is comprised of
microprocessor.
18. A method for producing an aortic pulse waveform, defined by
measurement of one of blood pressure, blood flow, or blood
constituent concentration, comprising the processing steps of:
measuring simultaneously of an aortic pulse waveform and an
arterial pulse waveform; calculating of a full aortic pulse
waveform, defined as the aortic pulse waveform without the effect
of atrial diastolic blood flow demand, and a full arterial pulse
waveform, respectively, defined as the arterial pulse waveform
without the effect of atrial diastolic blood flow demand;
calculating of an aortic arterio-venous pulse waveform, defined by
subtraction of measured aortic pulse waveform from said full aortic
pulse waveform, and an arterial arterio-venous pulse waveform,
defined by subtraction of measured arterial pulse waveform from
said full arterial pulse waveform; utilizing a system
identification to produce a state-space linear model to define a
transfer function relationship between the full arterial pulse
waveform as an input and the full aortic pulse waveform as an
output; utilizing the transfer function relationship with
continuous input full arterial pulse waveform to estimate the
output full aortic pulse waveform; utilizing a system
identification to produce a state-space linear model to define a
transfer function relationship between the arterial arterio-venous
pulse waveform as an input and the aortic arterio-venous pulse
waveform as an output; utilizing the transfer function relationship
with continuous input arterial arterio-venous pulse waveform to
estimate the output aortic arterio-venous pulse waveform; and
estimating of the aortic pulse waveform by subtracting the
estimated aortic arterio-venous pulse waveform from the estimated
full aortic pulse waveform.
19. The method claim of 18, wherein the aortic pulse is replaced
with a pulmonary artery pulse and the said peripheral arterial
pulse is replaced by a peripheral pulmonary arterial branch
pulse.
20. A method for producing an aortic pulse waveform as claimed in
claim 18, further comprising the steps of providing a processor
unit and configuring the processor unit to perform all said
processing steps.
21. A method for determining cardiac performance parameters as
claimed in claim 20, wherein said processor unit is comprised of
microprocessor.
22. A method for estimating cardiac output of the left or right
ventricular chamber using measured cardiac output from a right or
left ventricular chamber, respectively, adjusted using a balance
parameter; the cardiac output being adjusted using one of a
Left-Right Balance (LRB) parameter and a Right-Left Balance (RLB)
parameter.
23. The method claim of claim 22, wherein said estimating of the
left ventricle's cardiac output (LCO) is equal to the right
ventricle's cardiac output (RCO) divided by (1-the Right-Left
Balance (RLB) parameter.
24. The method claim of claim 22, wherein said estimating of the
right ventricle's cardiac output (RCO) is equal to the left
ventricle's cardiac output (LCO) multiplied by the Left-Right
Balance (LRB) parameter.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application is a non-provisional application which
claims the benefits of provisional application Ser. No. 61/700,892
filed on Sep. 14, 2012.
BACKGROUND OF THE INVENTION
[0002] 1. Technical Field
[0003] This invention relates generally to methods, devices, and
systems used in the field of cardiovascular medicine for clinically
assessing of a patient's health condition. More particularly, it
relates to a new and improved system and method for monitoring
cardiac blood flow balance between the right and left heart
chambers based on arterial pulse waveforms representing blood
pressure or flow or constituent concentration per cardiac
cycle.
[0004] The description below is merely provided for general
background information and for assisting in understanding the
technical field to which the present invention is related. As is
generally known in the art, the circulatory system in human beings
is responsible for transporting oxygen and other nutrients to the
cells of the human body. The circulatory system includes a heart,
arteries, capillaries, and veins. In a healthy patient, the heart
pumps blood with a certain pressure and volume so as to ensure that
proper blood circulation is realized within the human body.
Therefore, a discussion of the blood flow in the heart and
congestive heart failure (CHF) will now be provided initially.
[0005] With attention directed to FIG. 1 of the drawings which is
labeled "Prior Art", there is illustrated a diagram of the
cardio-vascular system and its cycle. The heart 2 consists of four
chambers; namely, the right atria 4, the right ventricle 6, the
left atria 8 and the left ventricle 10. The right atria 4 receives
carbon dioxide-filled blood which is returning from the body
through the superior vena cava 12 and the inferior vena cava 14
where they join in the central venous. The right ventricle 6
receives the blood from the right atria 4 via a tricuspid valve 16
located between the two chambers. The right ventricle 6 pumps the
blood via a pulmonary valve 22 into the pulmonary artery 18 which
branches into the right pulmonary artery 18(a) for carrying the
blood to the right lung 20(a), and the left pulmonary artery 18(b)
for carrying the blood to the left lung 20(b). After receiving
oxygen in the lungs 20(a) and 20(b), the oxygen-filled blood is
returned to the left atria 8 of the heart 2 via the right pulmonary
vein 24(a) and the left pulmonary vein 24(b). The blood is then
passed into the left ventricle 10 via mitral valve 26. Next, the
left ventricle 10 pumps the blood via the aortic valve 30 into the
aorta 28 which branches into the ascending aorta for delivery
throughout the upper body 31(a) heads and arms, and the descending
aorta 31(b) for delivery throughout the lower body including the
trunk and legs via the network of arteries, capillaries, and
finally returned to the heart 2 via the superior vena cava 12 and
the inferior vena cava 14, which both merge into the center venous
32. Both the right atria 4 and the left atria 8 pump
simultaneously, and both the right ventricle 6 and the left
ventricle 10 pump simultaneously. In every pumping cycle, each
chamber undergoes an expansion cycle called diastole followed by a
contraction cycle called systole. On the ECG waveform, the Atrial
diastolic phase extends between the S in the QRS complex until the
start of the following P wave, while atrial systole is between the
start of the P wave and extends through the following S in the QRS
complex. Similarly, the Ventricular systolic phase is between the R
in QRS-complex and the following end of the T-wave, while
ventricular diastolic phase is between end of the T-wave and the
following R in the QRS-complex. The atrial systolic phase occurs
mostly during ventricular diastolic phase with a small fractional
overlap to ventricular systole, while the atrial diastolic phase
starts with the ventricular systolic phase but also extends and
overlaps a substantial portion of the ventricular diastolic phase.
Atrial diastolic phase has a longer duty cycle as compared to
Atrial systolic phase.
[0006] The left ventricle (LV) and the right atrium (RA) are
connected in a supply and demand relationship. Also the right
ventricle (RV) and the left atrium (LA) are connected in a supply
and demand relationship. A balance therefore must be achieved for
each supply and demand relationship between the LV and RA (termed
Left Right Balance (LRB)), and between the RV and LA (termed Right
Left Balance (RLB)). Similarly a balance must be achieved across
the two pairs of supply and demand relationships (LV,RA) and
(RV,LA). Such a balance of the LRB and RLB balance parameters is
termed the Systemic-Pulmonary circulation balance parameter (SPB)
indicating a shift of the blood supply towards the systemic portion
of the circulatory cycle or the pulmonary portion of the
circulatory cycle. When the SPB is normal, this can be used to
validate assumptions of the Fick equation that the left and right
ventricular cardiac output on average are equivalent, as presumed
in healthy individuals.
[0007] In healthy human beings, the right heart's cardiac output
(RCO) blood flow is presumed equivalent to the left heart's cardiac
output (LCO), on average, since the cycle must maintain steady
blood flow through out the circulatory system. However, this is not
true in patients with certain types of heart disease. For example,
an infarction or arrhythmia on one side of the heart may lead to
insufficiency in balance between the RCO and LCO. Thus, cardiac
output (CO) is an important indicator not only for the diagnosis of
the certain types of heart disease, but also for the continuous
monitoring of a patient's health condition.
[0008] One basis for most common cardiac output-measurement systems
is given by the well-known equation CO=HR.times.SV, wherein CO is
the cardiac output defined as blood flow in liters/min and is equal
to the heart rate (HR) times the stroke volume (SV). The stroke
volume is usually measured in liters, and the heart rate is usually
measured in beats per minute. This equation explains that the
amount of blood the heart pumps out over a unit of time (e.g., a
minute) is simply equal to the amount it pumps out on every beat
(stroke) times the number of beats per time unit.
[0009] If the right heart's cardiac output RCO (blood flow out of
the right ventricle into the pulmonary artery) mostly exceeds the
left heart's cardiac output LCO (blood flow out of the left
ventricle into the aorta) then blood accumulation in the pulmonary
system (lungs) is evident, and which may lead to congestive heart
failure (CHF). Similarly, if the left heart's cardiac output LCO
mostly exceeds the right heart's cardiac output RCO then blood
accumulation is evident in the body's vascular system and tissue
fluids will accumulate leading to inflation in tissue or vascular
diseases such as high blood pressure, and potentially may play a
factor in the development of varicose veins. Therefore, a measure
of the balance between the right heart's and left heart's cardiac
outputs is necessary for monitoring the overall health of the
cardiovascular and pulmonary systems.
[0010] In the inventor's view, the cardiac system is basically a
control system with two primary objectives. The first primary
objective is to maintain oxygen delivery to the tissue via adequate
perfusion (blood flow and blood pressure), which is accomplished by
varying the cardiac parameters such as heart rate, stroke volume,
and contractility, and blood pressure. The heart's second primary
objective is to maintain a target blood flow balance between the
right and left heart chambers.
[0011] While the cardiac output balance between the right heart and
the left heart chambers is an important clinical parameter for the
purposes of diagnosing, treating, and monitoring cardiovascular
disease, the conventional prior art techniques, such as
thermodilution or dye dilution, for measuring this balance however
requires intravascular or intramyocardial instrumentation which
carries significant risks to the patient that includes increased
perioperative morbidity and mortality, and increased long-term
risks, such as stroke and pulmonary embolism. Additionally,
intravascular instrumentation can only be performed by highly
trained specialist which severely limits the availability of
qualified physicians capable of implanting the device, thereby
increasing the cost of the procedure.
[0012] Also, the estimation of the balance between right heart's
and left heart's cardiac outputs by these conventional methods
would necessitate measurement of both left and right cardiac
outputs independently and then comparing (for example by
subtraction or ratio or so any other means of comparison) both to
obtain the measure of the balance. In addition, less invasive
methods are also known in the prior art, but they have proved to be
less accurate in their measurements. Further, the traditionally
non-invasive methods are not able to distinguish between the left
cardiac output and the right cardiac output.
[0013] 2. Prior Art
[0014] In U.S. Pat. No. 5,211,177 to Chesney et al., there is
disclosed a vascular impedance instrument which includes a
transducer to obtain a digitized arterial blood pressure waveform.
The digitized data is used to determine cardiac output and to
subsequently obtain measurements of impedance parameters using the
modified Windkessel model of the arterial system. The instrument is
used as an aid in diagnosing, treating and monitoring patients with
cardiovascular disease.
[0015] In U.S. Pat. No. 7,651,466 to Hatib, Feras et al., examines
the pulse contour of arterial pressure pulses to map using an
linear model (Windkessel method) into a flow pulse for purposes of
estimating cardiac output. This does not take into account the
relationships between the ventricular arterial supply and atrial
venous demand on the shape of the pressure waveform, but it
analyzes the segment component of waveform prior to the onset of
the interaction of the atrial venous demand on the ventricular
arterial supply. Calibration of the model compliance parameters are
required for estimating the dynamic relationship between blood
pressure and flow rate.
[0016] In U.S. Pat. No. 8,282,564 to Parlikar, et. al., Similarly
uses a lumped compliance model to model (Windkessel method) the
relationship between arterial pressure and flow rate to estimate
cardiac output, again the interaction effects between the
ventricular arterial supply and atrial venous demand are not
considered on the shape or morphology of the arterial pressure
waveform.
[0017] In EU Patent No. EP 1,848,330 B1 to Bennett, Tommy, et. al.
Uses right ventricular pressure (RVP) pulse contour to estimate the
stroke volume using a Windkessel pulse contour method. This method
uses the full dynamic waveform information in the RVP to relate to
the stroke volume information. It does not selectively choose
specific landmarks on the pressure waveform to relate to volume
information in a static method as in this invention.
[0018] In U.S. Pat. No. 5,368,040 to James Carney, describes
analysis of ventricular pressure waveform and derives landmarks
identified from first and second derivative inflection points or
zero-crossing points or as related to features on an ECG waveform.
It does not attempt to map these pressure values to volumetric
information for purposes of estimating cardiac volumes or flow
rates given pressure information.
[0019] The inventor of the instant invention, however, is not aware
of system for measuring of a cardiac blood flow parameter between
the left chamber of the heart and the right chamber of the heart
that uses a sensor device to measure one of blood pressure and
blood flow and blood constituent concentration of a patient so to
generate an arterial pulse signal, a process unit for generating a
full arterial pulse signal from the arterial pulse signal, the
processor unit generating an arterio-venous pulse signal by
subtracting the arterial pulse signal from the full arterial pulse
signal, and the processor unit generating the balance parameter by
calculating the ratio of the area under the curve of the
arterio-venous pulse signal, measured by integration over the
cardiac cycle, to the area under the curve of the full arterial
pulse signal, measured by integration over the same cardiac cycle.
The area under the curve can be calculated by the integration of
the pulse waveform using any integration method, preferably
trapezoidal approximation method. Whereas the DC offset can be
either included in the integration, or normalized, or removed by
subtraction. The use of the word pulse throughout herein is used to
more generally indicate the measured signal of type such as a blood
pressure or blood flow rate or blood constituent concentration
waveform.
[0020] Therefore, it would be desirable to provide a system and
method for measuring of a cardiac blood flow parameter between the
left chamber of the heart and the right chamber of the heart useful
in creating a clinical assessment of a patient's health condition.
Further, it would be expedient that the system of the present
invention be capable of performing as a clinically useful tool for
the purpose of diagnosing, treating and monitoring the performance
and function of a patient's liver, heart, lungs, brain, kidneys and
other organs of the body. The present invention represents a
significant improvement over the aforementioned prior art which is
hereby incorporated by reference in their entirety.
BRIEF SUMMARY OF THE INVENTION
[0021] In broadest terms, a new and improved method is provided for
the estimation of balance parameters between the right heart
cardiac output and the left heart cardiac output based upon the
analysis of the pulse wave of the arterial pulsation. The analysis
includes the use of morphologic features of the Dicrotic Notch. The
Dicrotic Notch is typically present throughout the Aortic arterial
vasculature tree which is used to derive the Left-Right Balance
(LRB) parameter. The Dicrotic Notch is also typically present
throughout the Pulmonary Artery vasculature tree which is used to
derive the Right-Left Balance (RLB) parameter.
[0022] The inventor has developed a new theory for explaining the
morphology of the pulse wave of the arterial pulsation (arterial
pulse pressure or flow waveforms) and specifically the occurrence
of the Dicrotic Notch feature residing therein. The present
invention is based upon the realization by the inventor that the
morphology of an arterial pulse waveform lies in the circulatory
nature of the blood in the vascular system. Most of the cardiac
monitoring publications heretofore only analyzed either the left
heart's performance or the right heart's performance without
considering the interaction effects therebetween. In other words,
the fact that the left and right heart cycles complement each other
to complete the circulatory system was largely ignored.
[0023] The relationships between the right heart's cycle and the
left heart's cycle cannot be ignored as there is a continuity of
mass flow between them. Therefore blood flow, and subsequently
characteristics of pressure behavior, on one side of the cycle must
affect the other side of the cycle. The vascular pathways of
arteries and veins are interconnected by the capillary bed which
forms the connectivity between the right and left heart chambers,
and establishes the interdependency and continuity between them in
terms of blood flow. While blood pressure between the right and
left heart chambers is not directly interdependent, because of the
capillary bed therebetween, the blood flow is directly
interdependent as it has to be continuous between the left and the
right heart chambers. This direct relationship in blood flow
results, as will be explained later, in an indirect blood pressure
interdependency consequently. Although blood pressure pulse is
related to blood flow pulse by an arterial compliance factor, a
balance parameter expressed as a ratio of two pressure components
will normalize out the effects of the compliance and therefore the
ratio will be equivalent, at least from a trend perspective, to a
balance parameter that is derived from a blood flow pulse
measurement. Similarly, the blood constituent concentration dynamic
variation on a per cardiac cycle basis is directly related to blood
flow as the concentration will change based on the speed of the
mass flow per unit time by the sensor. Therefore balance parameter
trends derived from blood constituent concentration will be
consistent in behavior to balance parameter trends derived from
blood flow.
[0024] Often the performance of the left side of the vascular
system is analyzed separately from the right side, which yields
weak conclusions that cannot be generalized. Traditional models
that failed to take into account the simultaneous effects of the
left and right heart cycles tend to bias the attribution of the
dynamics it models to one side over the other, and therefore fail
to represent the true dynamics fully. The venous side of the blood
system is acted upon by the right atrium (RA) pull the blood up
towards the heart. Therefore, the RA actually applies a pull force
(demand) on the venous blood flow. On the left heart's side, the
left ventricle (LV) actually applies a push force (supply) on the
arterial blood flow to force the blood through the capillary system
into the tissues.
[0025] These two events occur substantially in simultaneous timing
(in normal subjects) so that the left side's (LV) push is aided by
the right side's (RA) pull, in order to facilitate the movement of
blood throughout the capillary system in the tissues. The LV and
the RA are interconnected by the vascular tree (arteries, veins,
and capillaries). While the blood pressure in the venous side is
substantially directly independent of the blood pressure on the
arterial side as they are connected via the capillaries, the blood
flow between the arterial and venous side is substantially directly
dependent, as conservation of mass laws and continuity of flow must
apply. In other words, in healthy subjects, whatever blood mass
flow the venous side demands the arterial side must supply and vice
versa to complete the circulatory system. In subjects with
circulatory system disorders, in addition to blood in the vascular
system, some fluids are either stored or extracted from tissue
fluids or the interstitial space. Blood pressure is an equilibrium
point to the demand and supply of blood (or blood flow). It is
therefore important to view pressure in the context of such supply
and demand forces on blood flow. This invention exploits the
relationships of blood supply and demand to each other, and uses
them to derive information about the arterial supply of blood from
venous pulse information and about venous demand of blood from
arterial pulse information.
[0026] In view of the foregoing background, it is therefore an
object of the present invention to provide a system and method for
measuring of cardiac blood flow balance parameter between the left
chamber of the heart and the right chamber of the heart of improved
design and performance. It is another object of the present
invention to provide a system and method for clinically assessing
of a patient's health condition based upon a balance parameter
obtained by measured arterial pulse signal and assessing of the
patient's health condition based upon the morphology of the balance
signal. It is still another object of the present invention to
provide a system and method for generating a full arterial pulse
(FAP) signal, defined as the measured arterial pulse signal without
the effect of atrial diastolic blood flow demand factor. The full
arterial pulse (FAP) is defined by the measured arterial pulse
signal without the effect of atrial diastolic blood flow demand,
simulating an added component of arterio-venous (AV) pulse signal
so as to result in a full arterial pulse signal (i.e. representing
ventricular supply factor without atrial interaction effects). It
is still yet another object of the present invention to provide a
system and method for measuring of cardiac blood flow balance
parameter which utilizes system identification techniques for
modeling a transfer function relationship between a peripheral
artery AV pulse or FAP signal and an aortic AV pulse or FAP signal.
It is still yet another object of the present invention to provide
a system and method for measuring of cardiac blood flow balance
parameter which utilizes system identification techniques for
modeling a transfer function relationship between a pulmonary
artery peripheral branch AV pulse or FAP signal and the pulmonary
artery AV pulse or FAP signal.
[0027] These and other objects, features and advantages of the
invention are provided by a system for measuring of cardiac blood
flow balance between the left chamber of the heart and the right
chamber of the heart which includes a sensing device for measuring
one of blood pressure and blood flow and blood constituent
concentration of a patient so as to generate an arterial pulse
signal. A system identification technique is responsive to the
arterial pulse signal for generating a full arterial pulse
signal.
[0028] Advantageously, a processor unit is provided and is
configured to produce an arterio-venous pulse signal defined by
subtracting the arterial pulse signal from the full arterial pulse
signal. Further, the processor unit computes a balance parameter by
calculating the ratio of the areas under the curve, by integration
over the cardiac cycle, for the arterio-venous (AV) pulse signal to
the full arterial pulse (FAP) signal. Whereas the DC offset can be
either included, or normalized, or removed by subtraction. In a
preferred embodiment, a simple linear transfer function is
established between these two signals such as an auto-regressive
moving average model is computed using system identification
methods. In another preferred embodiment, a generalized state-space
linear model of the balance parameter relationship between these
two signals is computed using system identification methods.
[0029] In addition, a computational device generates a set of
physiological parameters from the balance parameter. A display
monitor receives the set of physiological parameters from the
computational device for displaying meaningful information to
indicate a clinical assessment of a patient's health condition.
[0030] These and other features and advantages of the disclosed
system for measuring of cardiac blood flow balance parameter reside
in the construction of parts and the combination thereof, the mode
of operation and use, as will become more apparent from the
following description, reference being made to the accompanying
drawings that form a part of this specification wherein like
reference characters designate corresponding parts in the several
views. The embodiments and features thereof are described and
illustrated in conjunction with systems, tools and methods which
are meant to exemplify and to illustrate, not being limiting in
scope.
BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWING
[0031] FIG. 1 is a simplified diagram of a patient's cardiovascular
system, useful in explaining the blood flow through the heart and
labeled "Prior Art";
[0032] FIG. 2 depicts the various waveforms of a measured arterial
pulse signal, a derived full arterial pulse signal, and a
calculated arterio-venous pulse signal. It also depicts the
simultaneously measured central venous pulse, and the systolic and
diastolic phases of the atria and ventricles, as well as the
delayed systolic and diastolic phases of the atria and ventricles
as measured at the radial artery;
[0033] FIG. 3 is schematic block diagram of a system for producing
a balance parameter and related information, constructed in
accordance with the principles of the present invention;
[0034] FIG. 4 is a system transfer function which illustrates the
time domain and frequency domain relationships between the input
and output time series;
[0035] FIG. 5 is flow diagram of an exemplary embodiment of the
software used by the microprocessor in FIG. 3 for generating the
balance parameter and related information;
[0036] FIG. 6 are waveforms useful in explaining an embodiment of
the invention on how cardiac volumetric and performance parameters
can be determined using aortic pressure measurements allowing
determination of the left ventricular pressure at certain key
pressure points, which can be useful in mapping pressure
information to volume information;
[0037] FIGS. 7(a) and 7(b) are useful in explaining an embodiment
of the invention on how to determine a cardiac auto-regulation rate
and a respiratory rate determined from the cyclic variability of
the balance parameter, computed for each cardiac cycle, and over a
long term series of cardiac pulses;
[0038] FIG. 8(a) depicts balance parameter waveforms useful in
depicting the respiratory cycle and the normal cardiac
auto-regulation function cycle, and 8(b) depicts disappearance of
cardiac auto-regulation due to abnormal heart conditions, whereby
its restoration is necessary for indicating restoration of normal
heart condition;
[0039] FIG. 9 is a flow chart of an exemplary embodiment of the
method for estimating aortic pulse waveform (AOP), defined by blood
pressure, blood flow rate, or blood constituent concentration,
based upon the estimated full aortic pulse and aortic
arterio-venous pulse;
[0040] FIGS. 10(a) and 10(b) is a flow chart of an exemplary
embodiment of the method of FIG. 6 used to determine ventricular
pressure static points of end-diastolic pressure (PED), and
pre-diastolic end-systolic pressure (PDPES) for use in computation
of ventricular end-diastolic volume, end-systolic volume. It also
describes the determination of PED and PDPES using aortic pressure
alone without ventricular pressure information.
[0041] FIG. 10(c) is a flow chart of an exemplary embodiment of the
present invention used to calculate ejection fraction (EF) and
cardiac output (CO) based upon PED and PDPES obtained from FIGS.
10(a) and 10(b);
[0042] FIG. 11 is a flow chart of an exemplary embodiment of the
method for estimating a second ventricular chamber's performance
based upon information from a first ventricular chamber;
[0043] FIG. 12 is a flow chart of an exemplary embodiment of the
method of FIG. 8 for using the balance parameter for measurement of
the cardiac regulation cycle;
DETAILED DESCRIPTION OF THE INVENTION
[0044] Before explaining the disclosed embodiments in detail, it is
to be distinctly understood at the outset that the present
invention shown in the drawings and predominantly described in
detail in association with a system for measuring of cardiac blood
flow balance parameter is not intended to serve as a limitation
upon the scope or teachings thereof, but is to be considered merely
for the purpose of convenience of illustration of one example of
its application. In particular, the present invention may be
applied to any other suitable physiological monitoring systems,
such as oximetry systems, ECG systems, and any other suitable for
monitoring of a patient's liver, heart, lungs, brain, kidneys and
other organs of the body.
[0045] Historically, the dicrotic notch was thought to represent a
reflection of a fast pressure pulse that is generated when the
aortic valve opens above its cracking pressure, and blood pressure
reflecting backward from the arterial side back toward the aorta.
When this delayed and damped pressure component of the pulse wave
overlaps with the original pressure wave component, then this would
result in the combined components creating the elevation called the
Dicrotic Notch. However, this was later largely dismissed and
replaced by the valve recoil effect theory. The valve recoil effect
theory hypothesized that the Dicrotic Notch was the result of the
valve recoil after its passive closure due to pressure drop in the
start of the ventricular diastole phase. The valve recoil was
believed to cause an elevation of the pressure post the initial
decrease in pressure levels.
[0046] These theories however fail to explain why in some cases the
Dicrotic Notch was completely missing in some patients, and in some
other patients two dicrotic notches would appear. The fact that the
Dicrotic Notch was not consistently seen in all subjects is another
indication that this feature could not be entirely attributed to
valve closure alone. The simultaneity of the valve closure event
with the drop in pressure at the aorta and such event creating a
change in the rate of reduction of blood pressure represented in
the occurrence of the Dicrotic Notch has lead to the false
presumption that the Dicrotic Notch was formed due to that valve
closure event alone. The inventor believes that the circulatory
cycle must be analyzed completely to consider interdependencies
between the arterial and venous side in terms of blood flow
continuity and timing effects.
[0047] It has been observed by the inventor that the atrial
diastolic blood flow demand on venous blood has an interaction
effect on the ventricular systolic arterial supply of blood such
that this interaction affects the blood flow and subsequently blood
pressure. Such interaction and timing characteristics between the
ventricular supply and atrial demand of blood is what primarily
defines the dicrotic notch feature of the aortic and arterial
pulse. When ventricles start systolic contraction to eject (supply)
blood, the atria start diastolic expansion to dilate to fill with
(demand) blood, however, the atrial dilation extends longer in
duration to also overlap with ventricular diastolic expansion.
During this phase and in the absence of further blood supply the
level of steep decline of the pressure is due to atrial diastolic
blood flow demand, which starts from demand zero level reaches a
maximum demand peak (i.e. supply trough) and then reduces to demand
zero level. When atrial diastolic blood flow demand ends, due to
atrial systolic phase beginning, then the decay rate of the
arterial side supply pressure returns to its natural decay rate due
to ventricular supply pulse pressure reduction alone at its natural
decay rate. As the timing of this interaction and the level of
supply and demand vary with patient's age or health state, there is
sometimes observed the disappearance of the Dicrotic Notch or in
other conditions the appearance of two Dicrotic Notches. This blood
demand and supply interaction effect is realized across all the
capillary interconnections between the venous vascular system and
the arterial vascular system, and accumulatively the effect becomes
stronger and more easily observed at larger arteries, especially at
the aorta and pulmonary main arteries.
[0048] As was previously discussed, the pulse pressure at the aorta
and other arterial peripheral branches exhibit "a forced" or
"positively acted upon" loss of pressure that is subsequently
followed by a rise in pressure, when this effect ends, resulting in
the Dicrotic Notch. This is the primary effect creating the
Dicrotic Notch. In other words, the two events (i.e. blood pressure
drop followed by rise) create the feature of the Dicrotic Notch.
The rise in the pressure post the valve's passive closure can not
be fully attributed to the valve's recoil post the closure event
alone. When the valve's recoil occurs, it represents a
substantially smaller secondary pressure perturbation than is often
observed in the aortic valve's pressure wave about the Dicrotic
Notch defined by the primary effect.
[0049] The atrial diastolic blood flow demand (i.e. venous side
"pull") on the blood flow overlaps simultaneously with the
ventricular systolic supply (i.e. arterial side push pressure) of
the blood flow, during which the supply overwhelms the demand of
blood and the pressure rises. The atrial diastolic blood flow
demand then further overlaps with ventricular diastolic phase,
which stops the blood supply due to heart valve closing effect, and
allows the arterial pressure pulse to decay on its own natural
decay rate. However, because of the continuous interaction with the
atrial diastolic blood flow demand, which becomes the overwhelming
dominant factor that reduces the blood pressure more rapidly at a
forced decay rate that is steeper than the natural decay rate. This
continues until the atrial systolic event start point, at which
point the demand stops and the arterial decay rate returns to the
natural decay rate. This interaction effects between demand and
supply is substantially primarily responsible for the presence of
the Dicrotic Notch. Such interaction profile is also varied or is
affected by the blood viscosity or coagulation time, and can be
used to estimate blood coagulation or coagulation variation. The
venous side demand for blood flow overlapping with ventricular
supply forces an increased blood flow through the capillaries
greater than the rate it normally would occur at due to the
arterial supply alone. It also lowers the realized pressure at the
capillaries from the pressure levels that would be realized to
accomplish the same flow rates due to arterial supply alone. This
increased flow rate and at reduced pressure levels enhances the
ability of blood cells to flow through the capillary bed more
easily. In other words, the atrial diastolic blood flow demand
assists the ventricular systolic blood flow supply in flowing the
blood through the capillaries. The arterial pressure is affected by
the presence or absence or interaction timing of the ventricular
supply and atrial demand. The arterial supply pressure is lowered
in the presence of the atrial (venous-side) demand pressure pulse.
This venous side demand represented by the atrial diastolic blood
flow demand pulse results in a reduction in the pressure on the
arterial side at a faster rate (forced decay rate) than normal as
compared to the case where the venous demand pulse was not present.
Increased flow rate (due to venous side demand) also reduces the
observed concentration at a fixed measurement point (per unit time)
of blood constituents, such as oxygen content. Similarly, blood
flow rate sensors also depict the dicrotic notch feature due to
blood flow acceleration in the presence of atrial demand and the
interaction of ventricular supply as discussed previously.
[0050] As defined herein, the term "arterio-venous pulse" or (AV
pulse) refers to the portion of the arterial pulse (representing
blood pressure or flow or constituent concentration) that depicts
level reduction due to the atrial diastolic blood flow demand for
blood (from the venous side), and the arterio-venous pulse is
therefore measured on the arterial side and not the venous side.
When the arterio-venous pulse demand or "pull" reaches its maximum,
the arterial pulse supply or "push" reduces (dips) to a low level
presenting the valley of the Dicrotic Notch, this forced decay rate
of the arterial pulse pressure or flow is restored back to its
natural (unforced) decay rate once the arterio-venous pulse is
completed representing the end of venous demand and the remaining
reduction rate of the arterial pulse pressure or flow occurs at its
normal decay rate. The subsequent cessation or stopping of venous
side demand for blood flow results in the arterial pressure
increasing back to what it would have been normally at (unforced
natural decay rate) due to reduced flow in the arterial side. The
Dicrotic Notch event itself is a representation of the
arterio-venous pulse occurrence. The arterio-venous pulse demand
"pull" accelerates the flow of blood by assisting the arterial
pulse supply "push".
[0051] Venous side pulsations simultaneously add a blood flow
component (i.e. increase blood flow) due to its demand for blood,
and force the blood to flow faster. The presence of the
arterio-venous demand pulse reduces the overall pressure to flow
the blood through the capillary channels from levels required to
achieve the same flow rates in the absence of the arterio-venous
demand pulse. This venous pulsation (blood flow pull or demand) is
therefore responsible (primarily) for the creation of the arterial
side Dicrotic Notch. Aortic valve recoil has a similar but smaller
(secondary) effect on the aortic pressure and arterial pressure.
Those two factors (venous demand, and aortic valve recoil) forming
the dicrotic notch are discernible. The difference in the effects
of the aortic valve recoil versus the venous side demand on the
dicrotic notch can become more obvious or separable in the presence
of cardiac abnormalities causing atrial filling timing to be not
synchronized with ventricular ejection. In these cases, multiple
notches can be observed along the dicrotic limb of the pressure
wave.
[0052] Valve characteristics do not change significantly from heart
beat cycle to cycle, but the venous side demand (due to atrial
filling) is variable from heart beat cycle to cycle. Therefore, the
variations in the arterio-venous pulse across heart cycle-to-cycle
can be safely attributed to variations in venous side demand and
arterial side supply. Because pressure measurements are used to
quantify the arterio-venous pulse and pressure is measured on the
arterial side, the area under the curve, by integration over the
cardiac cycle, for the arterio-venous pulse is expected to be
smaller than that for the arterial pulse itself as the capillary
bed small channels exist between the arterial and the venous sides.
On the other hand, if flow rate were measured on the arterial side,
then the flow out of the arterial side must substantially equal the
flow into the venous side, less a small component of fluid
potentially transferred in/out of the tissue.
[0053] Therefore, an arterio-venous pulse measured from a blood
flow waveform (or equivalent indirect methods such as blood
constituent concentration variation related to flow) that is
detected at the arterial side is expected to have an area under the
curve that is substantially equivalent to the area under the curve
for the venous flow waveform.
[0054] As a result of the foregoing discussion, it can be seen that
the fundamental basis of the present invention is founded upon the
development of new and novel theory for explaining the occurrence
of the Dicrotic Notch. The present invention describes novel
measures and physiologic parameters that are used as clinical
indicators of cardiovascular and pulmonary performance and health
status from the analysis of the pulse wave and its features,
including the Dicrotic Notch. Furthermore, this invention describes
methods for measurement of such novel parameters via a variety of
sensory devices and algorithms or processing steps. In addition,
this invention describes several utility applications for use of
the derived novel indicator into clinical applications.
[0055] Referring now in detail to the various views of the
drawings, there is illustrated in FIG. 2 a plurality of timing
waveforms which are useful in explaining the present invention. The
waveform A is the measured arterial pulse signal in a normal,
healthy patient. The waveform B is the estimated full arterial
pulse (FAP) signal representing the arterial pulse in the
hypothetical absence of venous (atrial) demand. The estimated full
arterial pulse (FAP) signal is based on a linear line interpolation
between the primary peak and the secondary peak of the arterial
pulse signal A. The waveform C is the derived or calculated
arterio-venous pulse (AV) signal obtained by subtracting the
observed or measured arterial pulse signal A from the full arterial
pulse signal B. The waveform D is the central venous pressure (CVP)
waveform. A delay (D1) between features in the central venous
waveform and the radial artery waveform will be present and
substantially equivalent to the delay between the aortic waveform
and the radial artery waveform (D2), because of equivalent vascular
lengths. This delay can be measured between the start of
ventricular systole and the start of the cycle of increasing
arterial pressure in the radial artery. Ventricular systole begins
substantially synchronously with the R event of the QRS complex of
the electrocardiogram (ECG) or electrogram (EGM) waveform. This R
event is also synchronous with the point z of the central venous
waveform D.
[0056] Other features of central venous waveform D are indicated as
known in the art a-wave, c-wave, v-wave. Indicated on the FIG. 2 is
the period of the following phases: atrial diastole AD, atrial
systole AS, ventricular diastole VD, and ventricular systole VS. It
is important to note that the depicted indication of systolic and
diastolic phases is from a volumetric perspective and not from a
pressure perspective. The difference being the iso-volemic periods.
Also indicated is the delayed period of each of these phases when
shifted in time by the previously indicated delay D1, and are
indicated as follows: delayed atrial diastole ADd, delayed atrial
systole ASd, delayed ventricular diastole VDd, and delayed
ventricular systole VSd.
[0057] We notice that there is small overlap between (AS and VS)
until ventricular valve open, however, there is a larger overlap
between (AD and VS), (AS and VD), and (AD and VD). Similarly, there
is small overlap between the delayed version of these phases (ASd
and VSd) however, there is a larger overlap between (ADd and VSd),
(ASd and VDd), and (ADd and VDd).
[0058] We note that in the interaction phase of (VSd and ADd) the
ventricular supply overwhelms atrial demand and the pressure rises
until the end of VSd, at which point the radial artery pressure
reaches its maximum systolic value. Then the following phase is the
interaction of VDd and ADd, however ventricular demand is blocked
by the presence of valves, and the remaining effect on the radial
pressure is the delayed atrial diastolic ADd demand. The pressure
waveform depicts a natural slope of pressure decay once the pulse
moves past the sensor, however, because of the contribution of the
delayed atrial diastolic ADd demand, the radial pressure is forced
to decay at a higher rate, until the next phase of interaction
begins. Then the following phase is the interaction of VDd and ASd
where atrial demand ends and the only remaining element is the
natural decay rate of pressure, starting at the peak post the
dicrotic notch. This natural rate of decay continues until the
cycle restarts again with the interaction of the VSd and ADd. The
period during the phase of interaction between VDd and ADd is the
period of arterio-venous pulse waveform C, when atrial diastolic
blood flow demand effect forces the slope of decay of the radial
pressure to be more negative than the natural rate of pulse decay
during the ventricular diastolic phase as indicated during the
interaction of the phases VDd and ASd.
[0059] Measurements obtained at the central venous or pulmonary
vein are generally low pressure about 8-12 mmHg but are higher
power and higher in signal to noise ratio than those obtained at
smaller veins. The central venous pulse is also significantly
modulated by the respiratory cycle.
[0060] While it is obvious to the reader that the calculation of
the balance parameter can be performed either from the arterial or
venous measurements, and yielding equivalent results of balance
parameter detection and trending, the remaining discussion herein
will focus on arterial side measurements because of their higher
measurement power and resolution, and easier interpretation of the
full arterial pulse and the arterio-venous pulse relative to
equivalent information from the central venous pulse.
[0061] The Dicrotic Notch is present in the waveform A of the
arterial pulse signal. The Dicrotic Notch is defined by a
characteristic valley which resides between the peak (point 1) of
the arterio-venous pulse (also the beginning of the diastolic phase
of the ventricle), which is also followed by a secondary peak
(point 3) which also marks the end of the arterio-venous pulse.
[0062] In the hypothetical absence of atrial (venous side) demand
for blood flow and if the arterial blood flow was solely due to
ventricular supply, the decrease from the peak (point P1) of the
arterial pulse wave A (representing blood pressure pulse or blood
flow pulse) is modeled and presumed to be substantially linear over
time, for simplification, until it reaches the start of the
settling phase. The settling phase of the pulse wave A is defined
to start at the secondary peak (point P3) after the Dicrotic Notch
valley, which also indicates the end of the arterio-venous pulse
and the secondary valley (point P5) of the pulse wave which defines
its end (end of the diastolic phase) and the beginning of a new
pulse wave (start of the systolic phase).
[0063] The delays between the CVP waveform D and the arterio-venous
waveform C observed at the radial artery are substantially
equivalent to the delays between the aortic pressure and the radial
pressure. It will also be noted of the symmetry in the shape of the
AV waveform C obtained by subtracting the arterial pulse waveform A
from the FAP waveform B.
[0064] In FIG. 3, there is shown a schematic block diagram of a
system 32 for producing a balance parameter for use in one
application, constructed in accordance with the principles of the
present invention. The system 32 includes a sensing device or
transducer 34, an A/D converter 36, a processor unit 38, a RAM 40,
a ROM 42, and a storage device 44. The sensing device 34 is
preferably a non-invasive arterial blood pressure or blood flow
waveform measurement device operatively connected to a patient 33
for generating an arterial pulse signal. This arterial pulse signal
is similar to waveform A in FIG. 2 and is fed as an input to the
processor unit 38 via the A/D converter 36.
[0065] Further, the system 32 of the present invention includes a
user input device 46 for inputting information to the processor
unit, such as a keyboard or input terminal and a display station 48
for displaying useful information. The storage device can be any
computer-readable media capable of storing information that can be
interpreted by the processor unit, such as a hard disc, floppy
disc, or other digital storage apparatus.
[0066] The sampling rate of the A/D converter 36 is set to at least
two times according to the Nyquist criteria, but is preferably set
to five times or higher than the highest frequency component of
interest in the pulse waveform of arterial blood pressure, or blood
flow signal, or blood constituent concentration so as to
satisfactory capture the same. The A/D converter is used to
digitized the analog arterial blood pressure or blood flow, or
blood constituent concentration waveform signal generated by the
sensing device. Alternatively, the processor unit could obtain the
blood pressure or blood flow, or blood constituent concentration
measurements in digital form directly from the sensing device so as
to avoid the need of the A/D converter.
[0067] The sensing device 34 is preferably a non-invasive blood
pressure, or blood flow, or blood constituent concentration
waveform measurement device. A non-invasive blood pressure sensor
such as a finger-cuff transducer unit using a counter pulsation
technique in which the waveform is detected by balancing the air
pressure in the finger cuff with the blood pressure in the
patient's finger. A finger-cuff transducer unit of this type is
commercially available from Ohmeda Monitoring Systems of Englewood,
Colo. Also, tonometric blood pressure methods that press on the
radial artery against the bone structure may be used to perform
blood pressure measurements. A tonometric blood pressure of this
type is the Vasotrac which is commercially available from Medwave.
However, there exist many other prior art sensor types that can be
used to measure either arterial and venous pulse pressure
waveforms.
[0068] For blood flow measurements, this can be achieved by using
non-invasive acoustic ultrasonic doppler frequency effect sensors,
or using an electromagnetic wave (such as light or RF), or by
directly using motion sensors, accelerometers, or using coriolis
sensors, or by thermodilution methods, or dilution of concentration
of a chemical such as lithium or dye or a radioactive material, or
dilution of concentration of a blood constituent or component, or
using differential pressure sensors measurements (for example
Venturi flow sensing or orifice pinhole type sensors). Similarly,
pulse flow can also be assessed using the dilution or concentration
of a radiotracer chemical element such as those used in nuclear
medicine, fluoroscopy, or in magnetic resonance functional
imaging.
[0069] For blood constituent concentration measurements, this can
be achieved by using a noninvasive optical plethysmography or using
fast response time electrochemical sensors that have specificity
and sensitivity to the target chemical analyte or blood constituent
component.
[0070] For blood pressure measurements, this can be obtain using
many sensor types including resistive or capacitive or inductive
measurement elements, and any of the following broad
categorizations of general types of sensing technologies including:
[0071] Invasive method include placement of a catheter to obtain
pressure measurements consisting of: [0072] Using implantable
sensors that reside in a vascular branch or in a cardiac chamber or
body organ. [0073] Using catheter mounted pressure sensors, or via
catheter fluidic channel pressure measurement over the extended
catheter using a pressure sensor that is either built into the
catheter or external to it. [0074] Minimally invasive methods
include access to the artery or vein channel using a cannula and
using a pressure sensor connected to the cannula channel to measure
the vascular artery or vein pressure. [0075] Non-invasive methods
including: [0076] Tonometric pressure pulse sensing methods [0077]
Optical plethysmography pulse sensing methods [0078] RF pulse
sensing methods [0079] Acoustic pulse sensing methods such as those
relating the Doppler frequency shift with the flow rate change.
[0080] Vibration pulse sensing methods such as placement of a
accelerometer or force sensor in proximity to the vascular channel
where pulse waveform or pulse velocity is intended to be measured.
[0081] Blood magneto-acoustic methods which use measurement of an
electric current that is resultant from ion movement in a magnetic
field.
[0082] Blood temperature dynamic variation can also be
representative of the dynamics of the blood pulse when measured at
a frequency high enough that is representative (at least twice as
high based upon the Nyquist criteria, but preferably five times as
high) of the dynamics of variations of the pulse. This is due to
the thermal dilution gradient of the blood as it is accelerated
past the sensor. This measurement can be also obtained more easily
non-invasively, for example, using infrared sensors.
[0083] Blood constituents dynamic variations are also indicative of
the pulsatile flow when measured at a frequency high enough that is
representative (at least twice as high based upon the Nyquist
criteria, but preferably five times as high) of the dynamics of
variations of the pulse as are presented in the arterial pulse.
[0084] The monitoring of the cardiac balance parameter in the
present invention can be accomplished using pulsations measured
from any artery, including, but not limited to, aortic, carotid,
pulmonary, femoral, radial, renal, and hepatic arteries. Similarly,
the monitoring of the cardiac balance parameter can be accomplished
using pulsations measured from any vein including, but not limited
to, central venous, superior vena cava, inferior vena cava,
jugular, pulmonary, femoral, radial, renal, and hepatic veins.
However, the signal power measured at a vein is much smaller than
an artery. Nonetheless, the same primary components related to
arterial supply and venous demand can be identified from venous
measurements per cardiac cycle of any of blood pressure, blood
flow, or blood constituent concentration and can be used for
quantifying the balance parameter.
[0085] A balance parameter determined by measurements from the
central venous, central venous branches, the aorta, or aortic
arterial branch is called Left Right Balance (LRB), and a balance
parameter determined from the pulmonary vein, pulmonary vein
branches, pulmonary artery, or pulmonary arterial branches is
called Right Left Balance (RLB).
[0086] In addition, the monitoring of the cardiac balance using a
pulse waveform measured from any artery or vein can be accomplished
with any of the aforementioned sensor types and sensory methods of
measurement including invasive, minimally invasive, or noninvasive
blood pulse monitoring, and including pulse waveforms that are
indicative of blood pressure, blood flow rate, or blood constituent
or components dynamic variation over time.
[0087] Under the control of a software program stored in the ROM 42
or the storage device 44, the processor unit 38 will derive the
full arterial pulse (FAP) signal, as shown in the waveform B in
FIG. 2, based upon the observed or measured arterial pulse signal
A. Then, the processor unit 38 will calculate the arterio-venous
(AV) pulse signal, as shown in the waveform C of FIG. 2, by
subtracting the measured arterial pulse signal from the FAP signal.
Finally, the processor unit will determine a balance parameter at
its output 49 of the present invention by calculating the ratio of
the area under the curve of the AV pulse signal, measured by
integration over the cardiac cycle, to the FAP signal, measured by
integration over the same cardiac cycle. This balance parameter 49
is sent by the processor unit to a computational device 50 for
generating a set of physiological parameters on its output 51 which
are useful to indicate a clinical assessment of the patient's
health condition.
[0088] The display station 48 receives the set of physiological
parameters 51 from the computational device 50 for providing visual
and audio monitoring of the patient's health condition and for
alerting medical personnel of changes in the patient's health
condition. As an option, a wired or wireless telecommunication
system 52 shown in phantom may be coupled to receive the set of
physiological parameters from the computational device 50 and for
transmitting wirelessly the set of physiological parameters to the
display station 48, which can be located at a remote site away from
the patient.
[0089] In the embodiment shown, the processor unit 38 is comprised
of a general-purpose microprocessor which is adapted to execute
software consisting of an operating system and one or more
applications, as part of performing the functions described
hereinbelow. Further, each of the various blocks illustrated in
FIG. 3 is merely exemplary in nature and is intended to show one
implementation of the system for providing the balance parameter
using the arterial pulse signal and should not be construed as
limiting. Indeed, the different embodiments may have widely varying
software programs, data structures, applications, processes, and
the like. Therefore, the various functions of each block may in
actual practice be combined, distributed, or otherwise arranged in
any suitable fashion.
[0090] The processor unit 38 includes a System ID Processing Device
39 for establishing a transfer function relationship between FAP
pulses of any artery in the left heart vasculature (as an input or
inputs to the model) and the aortic artery main pulse which
represents the primary source of such a pulse (as an output to the
model). Similarly, the System ID Processing Device 39 can be used
to establish a transfer function relationship between full arterial
pulses (FAP) or arterio-venous (AV) pulses of any two arteries,
defined in forward or backward input/output transfer function
modeling steps. For example between the aorta and any of its
arterial branches, or between the pulmonary artery and any of its
arterial branches. Such transfer function relationships relate the
pulses from a plurality of locations as input or output in order to
use a more readily observable signal to estimate a more difficult
to observe signal. For example, used in the right heart vasculature
tree (as an input or inputs to the model) and the pulmonary artery
main pulse as a primary source of such a pulse (as an output to the
model). Preferably, linear state-space models can be used for
establishing such transfer function relationship.
[0091] The aortic or pulmonary artery full arterial pulse (FAP) can
be used as the source (input) producing the (peripheral) full
arterial pulses (output) in the arterial branches, or the reverse
(or inverse) model can be used. Similarly, the aortic or pulmonary
artery arterio-venous pulse can be used as the source (input)
producing the (peripheral) arterio-venous pulses (output) in the
arterial branches, or the reverse (or inverse) model can be used.
Whereas the estimated output arterio-venous pulse can be subtracted
from the estimated output full arterial pulse to produce the
estimated observed arterial pulse.
[0092] Similarly, the System ID Processing Device 39 can be used to
establish a transfer function relationship between FAP pulses of
any artery and the arterio-venous (AV) pulse from that same pulse,
representing a model of the balance parameter. Preferably, a simple
ratio of the AV pulse to the FAP pulse, i.e. a linear relationship
of the first order can be determined. Also, a more generalized
linear model of any order identified using system identification
methods, preferably state-space model, can be used to describe the
balance parameter as a system relationship. Such a generalized
model will capture the relative amplitude ratio and timing delay
(phase shift) information between the AV and FAP signals.
Optionally, in a second step, given the balance parameter, or a
model of it, and the input FAP, then the output AV pulse can be
estimated.
[0093] Selection of either inputs or outputs allows for estimation
or prediction of an unknown output, given an identified model
parameters and measured inputs. A MISO (multiple inputs single
output) type model mapping the relationship between a plurality of
peripheral artery pulses (inputs, more easily measured) into the
main primary source aortic or pulmonary artery pulse (output, more
difficult to measure) for that vasculature is a preferred
embodiment for enhancing accuracy and enabling easier prediction of
the central pressure at the main vessel. For a single peripheral
pulse, a SISO (single in single output) model is preferably used.
Once a model is identified and the coefficients are defined, then
the model can be used with the plurality of input peripheral pulses
to estimate or predict the primary source arterial pulse.
[0094] Such stated pulse(s) can be produced by a measurement from
different types of sensors including measurement of blood pressure,
blood flow rate, blood temperature, or blood constituent
concentration, such as hemoglobin, or glucose, or oxygen, or carbon
dioxide content. The same principles of model identification can be
applied using either the raw measured pulses or the estimated FAP
pulses (full arterial pulse composed of the raw measured pulse plus
the arterio-venous pulse) using any of these types of
measurements.
[0095] Such models relating peripheral arterial (radial, femoral,
renal, etc.) measurements and central arterial (aortic or pulmonary
artery) measurements which can be identified using system
identification methods can be used for estimation of such unknown
blood constituent concentrations at a primary (central or main)
blood vessel using only the measured peripheral blood constituent
concentration. Such model based estimation methodology can lead to
more accurate systemic estimation/measurement of such constituent
concentration and therefore enable better medication delivery
titration. Such medication delivery titration is desired for
control or regulatory compensation of the desired constituent,
especially with closed-loop feedback control systems, of which
model free adaptive control or model predictive controllers (MPC)
are two preferred embodiments. Similarly, a linear system
identification, preferably state-space model, can be applied across
peripheral Full Arterial Pulses (FAPs) such that an estimation of
an output peripheral FAP from another input peripheral FAP can be
obtained.
[0096] While there are available many tools and system
identification strategies, the preferred embodiments of the present
invention utilizes a direct system identification which uses
state-space methods, and preferably in a multiple input single
output (MISO) or single input single output (SISO) configuration,
as more fully discussed below. The state-space MISO or SISO system
identification allows implementation of system identification
methods to the application of cardiovascular and pulmonary
physiologic parameter estimation, and in the further improvement of
application of the Balance parameter as input to such models to
yield enhanced estimations to those parameters as model outputs.
This invention applies the science of system identification to the
realization of a feasible cardiovascular and pulmonary physiologic
parameter estimation.
[0097] In FIG. 4, there is depicted a system transfer function G(t)
which illustrates the relationships in the time and frequency
domains between the inputs X(t) and outputs Y(t). As can be seen,
in the time domain the inputs X(t) are convoluted with the transfer
function G(t) to produce the outputs Y(t). On the other hand, in
the frequency domain the inputs X(f) are simply multiplied in order
to obtain the outputs Y(f). Thus, the linear state-space model
identification from frequency response function allows ease of
adaptation of the system model using a change to its model order to
optimize its approximation or estimation of the system outputs.
[0098] The state-space models (as described in reference 4, Linear
Estimation, Kailath, Sayed, Hassibi. Prentice Hall, 2000, which is
hereby incorporated by reference) are defined as
x(k+1)=Ax(k)+Bu(k)
y(k)=Cx(k)+Du(k)+n(k)
[0099] Where u(k), y(k), and x(k) are time series of real numbers
representing the input, output, and state variables, respectively,
of the system, and n(k) is time series of real numbers representing
the noise term which is assumed to be independent of the input
sequence u(k). A, B, C, and D indicate the coefficients vectors.
The state variables represent the unforced dynamics of the system
model.
[0100] By using the Fourier transform on both sides of both
equations of the state-space model, we obtain the following, where
.sup. indicates superscript, and exp indicates the exponent:
(exp.sup.jw)X(w)=AX(w)+BU(w)
Y(w)=CY(w)+DU(w)+N(w)
Where w is the frequency term, j indicates an imaginary number, and
Y(w), U(w), X(w), N(w) are the frequency transformed output, input,
noise, and state variables. The variables A, B, C, D indicate the
coefficients vectors. Where
G(exp.sup.jw)=G(z) at z=exp.sup.jw=D+C((zI-A).sup.-1)B at
z=exp.sup.jw and, Y(w)=G(exp.sup.jw)U(w)+N(w)
is the frequency response function (FRF) of the system. I
represents the identity matrix and .sup.-I represents a matrix
inverse.
[0101] It will be noted that many other conventional System
Identification methods exist and can provide substantially
equivalent implementations to the state-space methods preferred in
the present invention. These include (a) linear system
identification (SYSID) methods, (b) nonlinear system identification
methods, and (c) blind system identification methods. The
background of each is discussed in details with algorithms
describing prior art implementation details of such methods. The
methods described apply for SISO (single input single output), MISO
(multiple input single output), SIMO (single input (stimulus)
multiple outputs (responses)), and MIMO (multiple inputs (stimuli)
multiple outputs (responses)) configurations. This instant
invention can equivalently, without loss of generality, use any of
the other system identification methods mentioned below, and in any
configuration of the inputs and outputs mentioned previously.
[0102] For example, as described in the references 1 through 6
listed in the Appendix, which are incorporated herein by reference,
which describe linear systems estimation and system identification
methods, the linear SYSID parameteric and non-parameteric methods
include: [0103] AR [0104] ARX [0105] ARMA [0106] ARMAX [0107]
Generalized Linear [0108] Output Error [0109] Box-Jones [0110]
Continuous transfer function [0111] Discrete transfer function
[0112] Impulse realization [0113] User defined model [0114]
Principal components subspace identification [0115] Discrete
frequency transfer function from frequency response function [0116]
Continuous frequency transfer function from frequency response
function [0117] Maximum likelihood methods
[0118] The nonlinear SYSID methods include: [0119] Neural Networks
[0120] Fuzzy Logic [0121] Volterra Series [0122] Weiner models
(LMS, or recursive least square based) [0123] Wavelets analysis
[0124] Nonlinear state-space models
[0125] The blind system identification methods include: [0126]
Laguerre model based [0127] Deconvolution methods
[0128] The cardiovascular or pulmonary measured pulse waveform can
refer to any of the following in continuous waveform or discrete
values that are absolute or relative in value: [0129] Blood
pressure in any vascular location or cardiac chamber. [0130] Blood
flow rate in any vascular location. [0131] Blood volume in any
cardiac chamber. [0132] Blood constituent concentration or chemical
analyte concentration in any vascular location or cardiac
chamber.
[0133] The system identification of the present invention can be
either a SISO, if for example a single input channel was mapped to
a single output channel, or, preferably, a MISO if the identified
system was determined by mapping (or relating) multiple input
channels (stimuli, measured cardiovascular or pulmonary parameters)
into a single output channel (response, estimated cardiovascular or
pulmonary parameter). Multiple inputs mapping provides greater
informational content and therefore a better mapping accuracy to
the output.
[0134] Alternatively, a MIMO configuration can be used where the
input measured signals (measured cardiovascular or pulmonary
parameters) are used to calculate multiple output responses
(estimated cardiovascular or pulmonary parameters) in a single
application of system identification methods rather than multiple
applications of multiple subsystem identifications. Alternatively,
a SIMO configuration can be applied where a single (measured
cardiovascular or pulmonary parameter) is used as an input to
determine system relationship with multiple outputs (estimated
cardiovascular or pulmonary parameters).
[0135] FIG. 5 is flow diagram of an exemplary embodiment of the
software used by the microprocessor in FIG. 3 for generating the
balance parameter. The process 500 is performed by the processing
unit 38. Initially, at step 502 the measured or observed arterial
pulse waveform signal A of FIG. 2 is generated by the sensing
device 34 and is received by the processing unit 38. At step 504,
the processing unit calculates or derives the full arterial pulse
(FAP) waveform signal B of FIG. 2 based upon the observed or
measured arterial pulse waveform signal A. Next, in step 506 the
processing unit calculates the arterial-venous (AV) pulse waveform
signal C of FIG. 2 by subtracting the arterial pulse waveform
signal from the FAP waveform signal. Finally, at step 510 the
balance parameter 49 at the output of the processing unit is
calculated by the ratio of the area under the curve of the AV pulse
waveform signal, measured by integration over the cardiac cycle, to
the FAP waveform signal, measured by integration over the same
cardiac cycle, in step 508. Whereas the DC offset can be either
included in the integration, or normalized, or removed by
subtraction.
[0136] With attention directed to FIG. 6, there are shown a
plurality of waveforms which are useful in explaining an embodiment
of the present invention with respect to how certain key pressure
points on the left ventricular pressure (LVP) or right ventricular
pressure (RVP) signal can be determined based upon an aortic
pressure or pulmonary artery measurement (or an estimate thereof).
In particular, the key pressure points of interest include
end-systolic pressure (PES), pre-diastole end-systolic pressure
(PDPES), and end-diastolic pressure (PED). The timing timeline for
atrial and ventricular systolic and diastolic phases are depicted
to ease the interpretation of the full aortic pressure pulse, given
the context of previous discussions in FIG. 2. It is important to
note that the depicted indication of systolic and diastolic phases
is from a volumetric perspective and not from a pressure
perspective. The difference being the iso-volemic periods. The
atrial diastolic phase is 623. The atrial systolic phase is 624.
The ventricular diastolic phase is 621. The ventricular systolic
phase is 622. Adopting a pressure based convention for systolic and
diastolic phase definition simply modifies the start and end of the
defined phases.
[0137] PED substantially represents the pressure point on the
ventricular pressure signal occurring pre iso-volemic contraction
prior to the systolic contraction of the heart volume and after the
diastolic expansion phase end. The end-systolic pressure (PES) and
the PDPES pressure however are two separate pressure points on the
ventricular pressure signal. The PDPES is indicative of the
ventricular pressure value post the isovolemic relaxation phase and
prior to diastolic expansion. The PED is indicative of the
ventricular pressure value post diastolic expansion and
pre-isovolemic contraction phase. While PED detection is generally
easily detectable from the waveform of the ventricular pressure
signal directly, the PDPES pressure point is not detectable
directly therefrom since it is often affected by active forces of
chamber muscle compression changing its relative timing.
[0138] The atrial diastolic blood flow demand creating filling of
the atria and during the overlapping phase with ventricular
isovolemic relaxation (indicated as part of ventricular volumetric
diastole) is what creates the arterio-venous pulse, which is, as
described herein, detectable at the aorta. Therefore, the timing
signature marking the end of diastole for the atria must be
detectable in the aortic pressure at the end of the arterio-venous
pulse. This same event can be used to mark the end of the
isovolemic relaxation event for the ventricle, after the systolic
contraction phase end, at which point such pressure value in the
ventricle (PDPES) represents substantially the post iso-volemic
relaxation which occurs prior to diastolic expansion of the
ventricular volume.
[0139] As can be seen in FIG. 6, the first waveform 605 is the
measured ventricular pressure signal. The second waveform 601 is
the measured aortic pressure (AOP) signal. The third waveform 603
is the measured electrocardiogram (ECG) signal, but can be
equivalently replaced by an electrogram (EGM) signal. The fourth
waveform 604 is a representation of the volume of blood in the
ventricle. Finally, the fifth waveform 602 is the first derivative
of the aortic pressure (AOP) signal 601.
[0140] The first key pressure point is PED which is designated by
the dot 607 on the ventricular pressure signal 605. The second key
pressure point is PDPES which is designated by the dot 612 on the
ventricular pressure signal 605. The third key pressure point is
the maximum ventricular pressure which is designated by the dot 608
on the ventricular pressure signal 605. The fourth key pressure
point is the point where ventricular pressure is equal to aortic
pressure which is designated by the dot 609 located at a first
intersection of the ventricular pressure signal 605 and the AOP
signal 601. The fifth key pressure point is the point where the
ventricular pressure signal separates from and is no longer equal
to the AOP signal which is designated by the dot 610 located at a
second intersection of the ventricular pressure signal 605 and the
AOP signal 601.
[0141] Advantageously, the measurement of PDPES in this embodiment
of the present invention is obtained from the measured waveform of
the ventricular pressure signal 605 (or its estimate), and using
event timing information derived from the waveform of the aortic
pressure signal 601. The derivative 602 of the aortic pressure
signal depicts a small peak indicated by the dot 615, at which
point the ventricular chamber is at end systole, and the atrial
chamber is at end diastole. At that event, the atrial diastolic
blood flow demand ("pull effect") of blood interaction with the
ventricular blood flow supply ("push effect") of blood reaches the
phase end (end of arterio-venous pulse) where an aortic pressure
first derivative 602 inflection point peak at that point indicates
such event 615.
[0142] Therefore, the PDPES (the dot 612) from the ventricular
pressure 605 can be measured (or estimated) at that event's timing,
which is also indicated with an inflection point 615 in the first
derivative 602 of the aortic pressure 601. It should be noted that
the following observations can be made: [0143] 1. PED can also be
preferably detected at the ventricular pressure point when the ECG
or equivalently EGM waveform triggers its R event in the QRS wave
complex. PED alternatively can be measured from the ventricular
pressure waveform directly using its 1st derivative valley. [0144]
2. PES can be measured from the ventricular pressure waveform by
finding the maximum value. [0145] 3. The valve open event (the dot
609) is detected preferably by the maximum peak 619 of the aortic
pressure first derivative, or by the minimum valley of the aortic
pressure, or by a peak of the first derivative of the ventricular
pressure. [0146] 4. The valve close event (the dot 610) is detected
preferably by the minimum valley 620 of the aortic pressure first
derivative, or the valley of the aortic pressure at the dicrotic
notch, or a valley of the first derivative of the ventricular
pressure. [0147] 5. The duration of the ventricular iso-volemic
contraction event spans the duration of the end of P wave to R
event in the QRS complex. Alternatively, preferably, it can be
defined as spanning the QRS complex event duration. [0148] 6. The
duration of the ventricular iso-volemic relaxation event spans the
duration of the electrical silence period of the heart.
[0149] The Balance parameter can be assessed from the aortic artery
(or pulmonary artery) pulse signal using an estimation of the
arterio-venous (AV) pulse signal, defined as the subtraction of the
measured or observed arterial pulse from the full arterial pulse
(FAP) signal at these two locations of the aorta or pulmonary
artery. The aortic full arterial pulse (FAP_aortic) (or
equivalently the pulmonary artery full arterial pulse
FAP_pulmonary) is measured by extending a linear line 616 between
the maximum pressure (peak--dot 608) of the aortic pressure signal
601 until the end of the dicrotic notch, at which point it is
advantageously substantially synchronous in timing with the PDPES
(the dot 612), which is detected from the aortic pressure signal
(or its estimate) using a inflection peak (the dot 615) of the
first derivative of the aortic pressure signal as indicated in the
waveform 602 of FIG. 6.
[0150] Still referring to FIG. 6, the small peak 615 of the first
derivative (waveform 602) of the aortic pressure signal 601
indicated at the vertical line 613, corresponds with the
pre-diastole end systolic event (PDPES) which is indicated on the
ventricular pressure waveform 605 with the dot 612 and on the
aortic pressure waveform 601 with the dot 611. This PDPES event
also corresponds with the ventricular volume being the smallest and
is substantially coincident to the extended linear line 616 across
the aortic pressure signal 601 used to define the FAP_aortic pulse
606. The vertical line 613 intersects the first derivative
(waveform 602) of aortic pressure signal 601 at a inflection point
615.
[0151] Similarly, the end-diastolic event at which end-diastolic
pressure (PED) is detected, defined by the ventricular pressure
point post completion of the ventricular diastolic expansion, and
prior to the onset of isovolemic contraction, is indicated as
corresponding to the timing of the R-event 614 of the
electrocardiogram (ECG) (or equivalently of an electrogram (EGM))
QRS-complex in the ECG (or equivalently EGM) waveform 603. The PED
value on the ventricular pressure waveform 605 is indicated by the
dot 607. The maximum ventricular pressure is also indicated on the
waveform 605 by the dot 608, which is equivalently the maximum
aortic pressure on the waveform 601. The dot 609 indicates the
valve open pressure event, when both the ventricular pressure in
the waveform 605 and the aortic pressure in the waveform 601 become
equal. The dot 610 indicates the valve close pressure event, when
the aortic pressure in the waveform 601 departs from the
ventricular pressure in the waveform 605.
[0152] With attention directed to FIGS. 10(a) and 10 (b) of the
drawings, there are shown, when joined together, a flow chart 1000
of an exemplary embodiment of the unique method of the present
invention in FIG. 6 used to determine the PED and the PDPES values
on the ventricular pressure signal for assessing the cardiac
ventricular volumes and other cardiac performance parameters.
Initially, in step 1001 a measurement (or an estimate thereof) of
either the aortic pressure (AOP) waveform for the left heart
performance or the pulmonary artery pressure (PAP) for the right
heart is obtained, which is similar to the waveform 601 shown in
FIG. 6. A measurement of the electrocardiogram (ECG) (or
electrogram (EGM)) waveform is obtained in step 1002, which is
similar to the waveform 603 shown in FIG. 6. In step 1003, the
first derivative waveform of the waveform obtained from step 1001
is computed, which is again similar to the waveform 602 depicted in
FIG. 6. Next, in step 1004 the valve open event (the dot 609 in
FIG. 6) is determined by using the maximum peak of the first
derivative waveform obtained from the step 1003. Similarly, in step
1005 the valve close event (the dot 610 in FIG. 6) is determined by
using the minimum valley of the first derivative waveform obtained
from the step 1003.
[0153] Further, the end systolic pressure PES event is determined
in step 1006 by using the maximum value of the waveform obtained
from step 1001, which is similar to the dot 608 in FIG. 6. In step
1007, the full arterial pulse waveform of the aortic pressure AOP
(FAP_aortic) or the pulmonary artery pressure PAP (FAP_pulmonary)
from the step 1001 is calculated. In step 1008, the arterio-venous
(AV) waveform of the aortic pressure AOP (AV_aortic) or the
pulmonary artery pressure PAP (AV_pulmonary) from the step 1001 is
calculated. In step 1009, the end (the dot 611) of the
arterio-venous pulse event of the AV waveform in step 1008 is
determined. This AV pulse event becomes zero for that heart cycle,
which is substantially coincident in time with the pre-diastolic
end-systolic pressure PDPES event, which is the dot 612 in FIG. 6,
defined by the ventricular pressure point post completion of the
isovolemic relaxation and prior to the onset of diastolic
expansion.
[0154] Continuing in the process 1000, in step 1010 the Left-Right
Balance (LRB) or Right-Left Balance (RLB) parameter is calculated
dependent upon if AOP or PAP was used, respectively, in step 1001.
This Balance parameter is based upon the respective FAP and AV
generated using a transfer function, preferably using the ratio of
the AV area under the curve, measured by integration over the
cardiac cycle, to FAP area under the curve, measured by integration
over the same cardiac cycle, or by some other relating transfer
function, from the corresponding steps 1007 and 1008 as well a
measurement of the heart rate HR. Whereas the DC offset can be
either included, or normalized, or removed by subtraction. In step
1011, the timing of the R-wave event 614 in the QRS segment of the
ECG (or EGM) waveform 603, which is substantially coincident with
the end diastolic PED event (the dot 607), is determined as
illustrated in FIG. 6. The heart rate is also measured in the step
1011. In step 1012, the time difference between the R-wave event
614 in the step 1011 and the valve open event 609 in step 1004 is
calculated as a delta T1.
[0155] In step 1013, by using the slope of the line segment of the
aortic pressure (AOP) (or pulmonary artery pressure (PAP)) pressure
between the valve open event 609 and a short time after the valve
opening (up to a period of a point approximately mid-way 617 of the
maximum AOP (or PAP) pressure). The PED value of the ventricular
pressure signal 605 can be extrapolated backwards from the AOP (or
PAP) waveform signal 601 (or an estimate of it). The extrapolation
period is defined by calculating the time between the ECG R event
(614) and the valve open event (609) as delta T1 of step 1012. The
extrapolation slope is defined by measuring the linear line fit
across the AOP (or PAP) pressure points 609 and the selected point
617. This AOP (or PAP) line segment (x-values representing time,
and y-values representing pressure values) is extrapolated
backwards in time up to the x-axis point in time to estimate the
PED event (the dot 607), as defined by the time occurrence of the
R-event of the ECG (or EGM) waveform 603. The y-axis value based
upon the result of this extrapolation is the PED pressure value at
point 607, which is considered a first static pressure point. For a
tiny fraction of a second during the heart cycle, at this first
static pressure point, all positive and negative dynamic forces
applied on the contained blood volume in the chamber are nulled,
and the ventricle is in its largest volume. The first static
pressure therefore represents the pressure only due to the volume
of the blood in the ventricle at its largest volume.
[0156] In step 1014, the time difference between the end of the AV
pulse event in the step 1009 and the valve close event in step 1005
is calculated as a delta T2. Next, in step 1015 by using the slope
of the line segment of the AOP (or PAP) pressure between the valve
close event 610 and a short time before the valve closing (up to a
period of a point approximately mid-way 618 of the maximum AOP (or
PAP) pressure). The PDPES value of the ventricular pressure can be
extrapolated forward from the measured AOP (or PAP) pressure (or an
its estimate of it). The extrapolation period is defined by
calculating the time between the end of AV pulse event (611) and
the valve close event (610) as delta T2 of step 1014. The
extrapolation slope is defined by measuring the linear line fit
across the AOP (or PAP) pressure points 610 and the selected point
618. This AOP (or PAP) line segment (x-values representing time,
and y-values representing pressure values) is extrapolated forward
in time up to the x-axis point in time to estimate the PDPES event,
as defined by the time occurrence of the small peak 615 in the
first derivative (waveform 602) of the AOP (or PAP) pressure post
the valve close event. The y-axis value based upon the result of
this extrapolation is the PDPES pressure value at point 612, which
is considered a second static pressure point. For a tiny fraction
of a second during the heart cycle, at this second static pressure
point, all positive and negative dynamic forces applied on the
contained blood volume in the chamber are nulled, and the ventricle
is in its smallest volume. The second static pressure therefore
represents the pressure only due to the volume of the blood in the
ventricle at its smallest volume.
[0157] The inventor has discovered techniques on how to use PED and
PDPES for assessing cardiac ventricular end diastolic volume (VED)
and end systolic volume (VES), which can be further applied to
calculate ejection fraction and cardiac output, as will now be
explained in detail in connection with the process 1000a in FIG.
10(c). Since PED and PDPES represent two static pressure points
when the ventricular chamber is respectively the largest and
smallest, a linear relationship exists between the chamber volume
and chamber pressure at these two static pressure points.
Therefore, the ratio of the volume difference (VED-VES) to the
ratio of the pressure difference (PED-PDPES) at these two pressure
points, provides the linear scale ratio R (ml/mmHg) needed to
convert the static pressure points into volume. This volume
information only need to be obtained during an initial calibration
step, such as volume being measured by a noninvasive ultrasound.
Thereafter the calibrated ratio R can be used continuously to
obtain cardiac volume information using pressure measurements
alone, which can provide detailed information of cardiac volumetric
and performance parameters and their trends on a continuous cardiac
cycle per cycle basis.
[0158] Ejection fraction can then be computed from the normalized
ratio (VED-VES)/VED. Equivalently, in step 1016 of FIG. 10(c), the
ejection fraction (EF) is measured using only the pressure
parameters PED and PDPES, wherein EF=1-(PDPES/PED) is compute
advantageously without requiring any volumetric measures.
[0159] In step 1017, for single-time calibration purposes, the
ventricular volume is measured, and the end diastolic volume VED
and the end systolic volume VES is determined using conventional
methods such as ultrasound, X-ray, or dilution techniques while at
the same time estimating the PED and PDPES. In step 1018 the
patient-specific relatively-constant calibration ratio
R=(VED-VES)/(PED-PDPES) is then computed, wherein the variables
VES, VED, PED, PDPES are all measured or estimated for the same
cardiac chamber.
[0160] From this linear scale, the end diastolic volume VED=R*PED,
and VES=R*PDPES can then be continuously estimated at step 1019.
This method can be applied for any cardiac chamber. From cycle to
cycle, the VED, VES can change to maintain balance between the
right and left heart cardiac outputs, and therefore PED and PDPES
will also change.
[0161] Other known in the art derived clinical cardiac parameters
from VED and VES can also be obtained. In step 1020, this volume
difference (VED-VES) represents the stroke volume (SV) which can be
estimated using PDPES and PED from the process in FIGS. 10(a) and
10(b) described just above. Next, in step 1021 the cardiac output
(CO) is computed as CO=SV.times.Heart Rate. In step 1022 using CO
and Body Surface Area (BSA), the cardiac index (CI)=CO/BSA is
computed. In step (1023), the estimate of the mean arterial
pressure (MAP)=PED+(PES/3) is computed. In step 1024, the central
venous pressure (CVP) is measured or estimated.
[0162] Continuing with process, in step 1025, we compute systemic
vascular resistance (SVR)=(MAP-CVP)/CO is computed or is equal to
the difference in average pressure between arterial and venous
sides divided by the average flow rate. In step 1026, the stroke
work (SW) as SV.times.MAP is computed. In step (1027), the estimate
of contractility is computed using derivative of the ejection
fraction EF. In step (1028), the ratio R computed from one cardiac
chamber (e.g., the left ventricle) is used to compute the cardiac
performance parameters for other cardiac chambers (such as the
right ventricle) given its PED and PDPES pressure values.
[0163] In another aspect of the present invention, the inventor has
realized that the variability of the Balance parameter (LRB or RLB)
is a viable indicator of cardiac health state. The balance
parameter is a very sensitive and specific measure for pulse
variability, reflecting variation in supply and demand of blood,
rather than the variation in the blood pressure or flow alone. It
is also sensitive enough to detect respiration modulation effect of
the cardiac cycle and advantageously allows the detection of
respiration rate and parameters directly from the balance parameter
variations.
[0164] With reference now to FIGS. 7(a) and, 7(b), there are shown
a balance parameter indication with a plurality of auto-regulation
cycles modulating a plurality of respiratory cycles 701. Each point
on the plot is a cardiac cycle. The portion of the respiratory
cycles 701 between the points A and B in FIG. 7(a) are expanded and
enlarged in FIG. 7(b). The boxed area 702 in FIG. 7(b) is the
auto-regulation cycle, and the boxed area 703 is the respiratory
cycle. Thus, in this aspect of the present invention, the
respiration rate is detected from simple variation cycle of the
balance parameter. Furthermore, the Balance parameter is sensitive
enough to detect the cardiac auto-regulation cycle 702 which
maintains the regulation of the balance between the left and right
cardiac outputs. This longer period regulation cycle 702 was
previously not measurable or realizable, but is now advantageously
detected by the balance parameter of the instant invention.
Characterization of auto-regulation function is therefore an
important aspect of this invention. Detection of abnormalities in
the auto-regulation function and its cycle characteristics,
alarming against auto-regulation abnormalities, and display and
communication of auto-regulation information to users are all part
of the instant invention.
[0165] With reference to FIG. 8(a), there is depicted the normal
auto-regulation function to the cardiac cycle 805 as a vital
indicator to the health state of the heart. Biased balance
indication (i.e. too high, or too low) is an early indicator of
heart disease. In FIG. 8(b), there is also depicted a perturbation
at 806 causing the balance parameter to start an increasing trend
806 through 807, which is not be easily detectable using blood
pressure signal alone. As shown in FIG. 8(b), when the heart
experiences abnormal ischemia 811 the auto-regulation periodic
cycle pattern of the balance parameter disappears 813, and when
arrhythmia or fibrillation is experienced 812 a high variance of
the balance parameter is indicated. Abnormal cardiac conditions
such as arrhythmia or fibrillation are detected with the variance
of the balance parameter above a threshold and can be indicated for
an alarm. Restoration of auto-regulation periodic cycle is an
essential indication of restoration of normal function to cardiac
cycle. Absence of auto-regulation or other abnormal cardiac
conditions are detected with the variance of the balance parameter
below a threshold and can be indicated for an alarm. Abnormal
cardiac condition detection and indication can be used for closed
loop feedback in the delivery of therapeutic medications or to
invoke therapy using a pacemaker, cardioverter, or a
defibrillator.
[0166] With reference to FIG. 12, there is illustrated a flow chart
1200 of an exemplary embodiment of the method of FIGS. 7 and 8 for
using the balance parameter for measuring the cardiac regulation
cycle. By separating the respiration signal component 703 from the
regulation signal component 702, it is possible to obtain a better
assessment of each of the individual components independently.
[0167] In the process 1200, in the step 1201 an arterial pressure
waveform (i.e., from an peripheral artery or pulmonary artery or
aortic artery) is measured initially. In step 1202, a full arterial
pulse (FAP) and an arterio-venous (AV) pulse are calculated. In
step 1203, the balance parameter (LRB or RLB) is determined for
each of heart pulse over a long duration of the heart pulses. In
step 1204, there is applied optionally a filtering to smooth out
the balance parameter from step 1203. In step 1205, demodulation of
the respiratory signal carrier is applied to obtain the regulatory
signal carrier, or alternately demodulation of the regulatory
signal carrier is applied to obtain the respiratory signal
carrier.
[0168] However, this is difficult to achieve effectively since the
respiration rate is variable. One preferred method is to use
adaptive filters so as to filter out an undesired component (such
as either respiration or regulation) in order to obtain the desired
component (such as either regulation or respiration, respectively).
Adaptive filters require a reference signal to indicate the
component to be separated out. This reference signal is more easily
provided as the respiratory signal which can measured with a
variety of methods, including motion sensitive accelerometers, or
tension belt or modulation of the ECG waveforms, or nasal air flow
sensor. Also, there are conventional tools known in the prior art,
such as matrix-pencil methods for frequency estimation, can be used
to separate very effectively the component frequencies representing
the cardiac auto-regulation cycle from the respiration cycle.
[0169] In step 1206, a peak-valley detection is optionally applied
to the separated components from step 1205 to determine the
end-of-cycle period, frequency, and duty cycle. Spectral frequency
analysis (e.g. Fourier series) is performed in step 1207 to detect
the primary regulation and respiration rate. In step 1208, other
respiratory parameters can also be derived such as the inhalation
and exhalation duty cycle. The detected respiratory rate can be
utilized to monitor and provide alert for respiratory disorders,
including respiratory apnea in a highly sensitive, specific, and
reliable method, that is not affected by additional noise sounds
such as snoring or air flow measurement errors. This method can
also be combined with anesthesia ventilators in order to determine
patient restoration of respiratory function after respiratory
depression with anesthesia or pain medications, to allow
determination of optimal ventilator settings to wean patient of
ventilator and to control the ventilator in closed loop to achieve
patient weaning off ventilator. The respiration waveform can also
be used to more generally assess patient respiratory function and
health state. In step 1209, by using the auto-regulation cycle
rate, its period or frequency, or duty cycle, baseline trend,
maximum, minimum, mean and standard deviation statistical
distribution and properties, and heart rate can all be determined.
In step 1210, the auto-regulation cycle and respiratory cycle
characteristics and information can be used to assess patient
health state and alarm against abnormalities, and determine
therapeutic measures or decisions to be applied, including closed
loop control of medication delivery or cardiac interventional
devices. The balance parameter and auto-regulation performance
parameters can be used to optimize the treatment protocol of renal
dialysis such controlling the blood flow rate into the dialyzer to
avoid abnormal cardiac balance and autoregulation performance. The
balance parameter and auto-regulation performance parameters can
also be used in feedback closed loop control of dialysis devices to
achieve optimal dialysis treatment control. The balance parameter
and auto-regulation performance can also be used for detection and
alarming of cardiac stress, patient pain, or surgical nociception,
whereby marked trend variations in the balance levels and
regulation cycle are associated with elevated pain levels or
cardiac stress. The balance parameter and auto-regulation
performance parameters can be used to provide feedback control to
ventricular assist devices and counter-pulsation devices to help
optimize the blood flow for a more balanced auto-regulation of the
heart cycle. The balance parameter and auto-regulation performance
parameters can be also used for guidance and optimization of lead
placement locations of a pacemaker system to provide a more
balanced auto-regulation performance of the heart cycle. The
cardiac balance and autoregulation performance parameters can be
used to optimize blood flow between the mother and the fetus and
can be used to study the fetus cardiac regulation and balance as
well as the mother's cardiac regulation and balance, and the
interaction of both over the term of the pregnancy, to ensure that
both mother and fetus are receiving optimal blood flow, and to
ensure baby healthy rate of growth and nutrition with adequate
blood flow. Abnormal fetal blood flow in relationship to mother's
blood flow can be detected and alerted.
[0170] The variance (or similarly standard deviation) of the
Balance parameter has been advantageously found by the inventor to
be a strong indicator of the health status of the subject's cardiac
function. If the Balance parameter variance is too high as
illustrated at 812 in FIG. 8(b), then it is an indicator of cardiac
instability, potentially as an indicator or measure of surgical
nociception or pain, arrhythmia, ischemia or myocardial infarction,
as some causes for high balance variance. These are typically
associated with loss of the auto-regulation function indicated by
the balance parameter at 811 and 813. The absolute value of the
balance parameter is used as an indicator of cardiac imbalance
representing a biased performance toward the left or right heart
blood pumping (cardiac output). Such an imbalance can lead to the
accumulation of blood in parts of the circulatory system which can
lead eventually to cardiac arrhythmia, or heart failure such as
congestive heart failure.
[0171] If on the other hand the variance of the heart is too low
from its normal auto-regulation cycle variance baseline, then this
also is an indicator of hydration status or other cardiac problems
including, for example hypovolemia, or low hydration or low cardiac
output status, and is a strong predictor of heart failure. There
exists a margin of normal variance of the Balance parameter about
which the cardiac performance is automatically adaptively
maintained by the heart during the auto-regulation cycle.
[0172] It is still another aspect of this invention to use both the
Balance parameter as a mean target value set point and the Balance
parameter variance (or standard deviation) margin as a second
target set point in control of medication delivery that affect
these cardiac performance parameters to the desired set target
levels or margins. It is still yet another aspect of this invention
is to use the measurement of Balance parameter variance at 812 as
well as the absolute value of the Balance parameter at 807 for
purposes of monitoring, predicting, and/or alerting against cardiac
arrhythmias, fibrillation, ischemia, or heart failure such as
congestive heart failure (CHF), or detection of onset of seizures
and alarming against it.
[0173] In addition, valvular malfunction (regurgitation or
insufficiency) will affect the appearance or timing of the Dicrotic
Notch. It is thus envisioned that this invention can also use the
RLB Balance and/or LRB Balance parameters to diagnose the presence
or absence of as well as the degree or status of valvular
regurgitation or insufficiency. Furthermore, another aspect of this
invention is to use the balance parameter for indication of valve
regurgitation or insufficiency.
[0174] Further, a cardiovascular control support system that
provides pumping of a balloon in order to displace blood volume
whereby the device is placed in a vascular artery or vein is used
to achieve therapeutic or optimal Right-Left Balance levels.
Another aspect of this invention is to use the Balance parameter as
a process variable in a closed loop control system for controlling
a counter pulsation (balloon) apparatus in order to achieve a
desired set target value for cardiac parameters, including Balance
parameter, blood pressure, and blood flow rate or cardiac output. A
model predictive controller (MPC) or a model free adaptive
controller can be preferably used for this purpose.
[0175] In yet another embodiment of the present invention, an
estimated aortic pressure waveform can be derived by first
estimating its components of FAP_aortic and AV_aortic and then
subtracting the estimated AV_aortic waveform from the estimated
FAP_aortic waveform. In FIG. 9, there is depicted a flow chart 900
of an exemplary technique for determining the estimated aortic
pulse waveform, defined by one of aortic blood pressure, blood flow
rate, and blood constituent concentration. Initially, at step 901
the aortic pulse is measured. Simultaneously, at step 904 the
(peripheral) arterial pulse (e.g., radial or femoral or carotid) is
measured, defined by one of arterial blood pressure, blood flow
rate, and blood constituent concentration. In step 902, the
aortic's full arterial pulse (FAP_aortic) waveform is computed as
defined herein. Similarly, the peripheral artery's full arterial
pulse (FAP_arterial) is computed in step 905. Further, the aortic
arterio-venous pulse (AV_aortic) is computed in step 903. The
(peripheral) arterial arterio-venous pulse (AV_arterial) is
computed in step 906.
[0176] In step 907, the relationship or transfer function between
the FAP_arterial as input and the FAP_aortic as output can be
defined using many modeling and adaptive methods, including a
system identification method, as mentioned herein, preferably using
linear system identification producing a state-space linear model
to define such transfer function. In one preferred embodiment, the
mapping between FAP_arterial and the FAP_aortic, the
transformation's transfer function (preferably state-space model's
coefficients) can be identified easily. Such model creates a
calibration transfer function when both FAP_arterial and FAP_aortic
are available. In step 909, such calibration transfer function can
then be used with input FAP_arterial to produce a more accurate
estimate of the output FAP_aortic. Similar or equivalent steps can
be taken as above with FAP_pulmonary derived from the pulmonary
artery pulse replacing the FAP_aortic, defined by one of pulmonary
artery's blood pressure, blood flow rate, and blood constituent
concentration.
[0177] Continuing in the process 900, in the step 908 the
(peripheral) arterial pulse (radial or femoral or carotid)
arterio-venous pulse (AV_arterial) as defined herein, is used to
estimate the aortic's arterio-venous pulse (AV_aortic) waveform as
defined above. The relationship or transfer function between the
AV_arterial as input and the AV_aortic as output can be defined
using many modeling and adaptive methods, including a system
identification method, as mentioned herein, preferably using linear
system identification producing a state-space linear model to
define such transfer function. In one preferred embodiment, the
mapping between AV_arterial and the AV_aortic, the transformation's
transfer function (preferably state-space model's coefficients) can
be identified easily. Such model creates a calibration transfer
function when both AV_arterial and AV_aortic are available. In step
910, such calibration transfer function can then be used with input
AV_arterial to produce a more accurate estimate of the output
AV_aortic. Similar or equivalent steps can be taken as above with
AV_pulmonary derived from the pulmonary artery pulse replacing the
AV_aortic. Finally, in step 911 an estimate of the aortic pulse
waveform can be derived by subtracting the estimated AV_aortic
waveform from the estimated FAP_aortic waveform.
[0178] With reference to FIG. 11, there is shown a flow chart 1100
of an exemplary embodiment for estimating the other ventricular
chamber's performance based upon information from the one
ventricular chamber. The left heart's (VED-VES) is estimated, on
average, using a Balance adjusted right cardiac output (RCO)
divided by the heart rate (HR). If measured cardiac output is
performed in the right heart, the volume difference
(VED-VES)=RCO/HR. The parameter RCO can be adjusted using the
Balance parameter value (Left-Right Balance (LRB) from any aortic
branch artery or Right-Left Balance (RLB) from any pulmonary branch
artery) for the estimation of the left heart's cardiac output
LCO.
[0179] In step 1102 of FIG. 11, the left ventricle's (LV) cardiac
output is calculated as either LCO=(RCO/LRB) or LCO=RCO/(1-RLB).
Similarly, in step 1101 the right ventricle's (RV) cardiac output
is calculated as either RCO=LCO*LRB or RCO=LCO*(1-RLB). Both LRB
and RLB are <1.0, but the LCO will be greater than the RCO, with
the difference being attributed to inefficiencies of transfer (e.g.
backflow portion or tissue fluidic uptake). A higher RLB is
associated with a lower LRB and is typically associated with left
heart abnormal function. Similarly, a lower RLB is associated with
a higher LRB and is typically associated with right heart abnormal
function.
[0180] Summarizing the above description of the present invention
with respect to FIGS. 9 and 10, a plurality of peripherally
measured arterial pulse in step 904 is used to derive a
FAP_arterial waveform in step 905, and a main artery (aortic or
pulmonary artery) measured arterial pulse in step 901 is used to
derive its FAP in step 902. Then, applying a calibration step, the
plurality of input derived (peripheral)-arterial waveform(s) is/are
mapped to output primary main artery FAP_aortic (or FAP_pulmonary)
using a transfer function defined by a system identification method
at respective steps 907 and 908. Preferably, the system
identification method is a linear state-space model in a MISO
(multi-input single-output) or SISO (single-input single-output)
configuration.
[0181] Then, in the steps 909 and 910, there is provided
continuously acquiring (peripheral) arterial pulse, deriving the
FAP_arterial and employing the previously identified transfer
function to produce an estimate of the FAP_aortic. In the step 911,
there is provided using the continuously produced estimate of the
FAP_aortic to estimate the left ventricular (LV) pressure points of
PED and PDPES obtained from the process 1000. Equivalently, there
is provided using the continuously produced estimate of the
FAP_pulmonary to estimate the right ventricular (RV) pressure
points of PED and PDPES.
[0182] In the process steps of FIG. 10(b), there is provided using
the PED and PDPES estimated in conjunction with measured or
estimated (adjusted) cardiac output and heart rate and Balance
parameter to produce cardiac parameters including, but not limited
to, absolute values of end-diastolic volume, end-systolic volume,
and additional cardiac parameters known-in-the-art such as stroke
volume, cardiac output, ejection fraction, systemic vascular
resistance, contractility, and stroke work. Furthermore, there is
produced an estimate or measurement of the duration of the
ventricular iso-volemic contraction period and the ventricular
iso-volemic relaxation period.
[0183] From the foregoing detailed description, it can thus be seen
that the present invention provides a system for measuring of
cardiac blood flow balance parameter between the right chamber of
the heart and the left chamber of the heart which includes a sensor
device for measuring one of blood pressure and blood flow rate and
blood constituent concentration of a patient so as to generate an
arterial pulse signal. A processing unit is responsive to the
arterial pulse signal for generating a full arterial pulse (FAP)
signal, an arterio-venous (AV) pulse signal, and a balance
parameter. A computational device is responsive to the balance
parameter for further generating a set of physiological parameters.
A display station device is responsive to the set of physiological
parameters from the computational device for displaying meaningful
information.
[0184] While there has been illustrated and described what is at
present considered to be a preferred embodiment of the present
invention, it will be understood by those skilled in the art that
various changes and modifications may be made, and equivalents may
be substituted for elements thereof without departing from the true
scope of the invention. In addition, many modifications may be made
to adapt a particular situation or material to the teachings of the
invention without departing from the central scope thereof.
Therefore, it is intended that this invention not be limited to the
particular embodiment disclosed as the best mode contemplated for
carrying out the invention, but that the invention will include all
embodiments falling within the scope of the appended claims.
APPENDIX: REFERENCES
[0185] 1. System Identification: Theory for the User, Lennart
Ljung, Prentice Hall, 2nd edition, 1999. [0186] 2. System
Identification: A Frequency Domain Approach. Rik Pintelon and Johan
Schoukens Wiley-IEEE Press, 1st edition, January, 2001 [0187] 3.
Blind Equalization and System Identification: Batch Processing
Algorithms, Performance and Applications, Springer, 1st Edition,
January 2006. [0188] 4. Linear estimation, Kailath, Sayed, Hassibi.
Prentice Hall, 2000. [0189] 5. Multivariable System Identification
For Process Control, Y. Zhu, Elsevier Science; 1 edition (October,
2001) [0190] 6. Modeling of Dynamic Systems, L. Ljung, Prentice
Hall; 1 edition, May 1994.
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