Method and apparatus for continuous measurement of cardiac output and SVR

A method of continuously monitoring cardiac output and SVR of a patient by analyzing the blood pressure signal. The signal can be measured directly by the use of catheters placed into various arteries, in particular a radial artery or a femoral artery, or non-invasive methods such as electrooptic means, or using a piezoelectric pressure transducer. Various parameters are extracted from the blood pressure signal waveform in order to calculate the cardiac output of the patient and other hemodynamic data.

BACKGROUND OF THE INVENTION 
1. Field of the Invention 
The present invention relates to a new method and apparatus for monitoring 
and/or measuring cardiac output and SVR of a patient by analyzing a blood 
pressure signal which contains various parameters related to particular 
characteristics of the patient's vascular system. 
2. Description of the Related Art 
Cardiac output is the ultimate expression of cardiovascular performance. 
The term "cardiac output" indicates the quantity of the blood ejected each 
minute by either the right or left ventricle. Absolute cardiac output in 
an in vivo pulsating flow system is the product of the stroke volume and 
the heart rate frequency per minute. Stroke volume quantitatively defines 
the beat-to-beat volumetric performance of the heart as an intermittent 
flow generator or pulsating pump. The performance of the pump on a 
beat-to-beat basis and thus the magnitude of stroke volume is determined 
by three major independent variables: pre-load, contractility and 
after-load. Each major determinant of stroke volume has its own subsets of 
major and minor independent variables. The integrated modulation of all 
variables contributes to the stroke volume generated by the ventricles for 
each heartbeat. The parameters associated with cardiac output are useful 
in that they can be employed to evaluate the overall cardiac status of 
critically ill patients, patients with suspected cardiovascular and 
pulmonary diseases, patients undergoing surgery, and any situations that 
require blood pressure monitoring. 
The radial stretch of the ascending aorta involved in the ejection of blood 
by the left ventricle initiates a pressure wave which propagates down the 
aorta and its various branches. The pressure wave travels with a finite 
velocity that is considerably faster than the actual forward movement of 
the blood itself, and is a wave which pulsates as it reaches the 
peripheral arteries. The velocity of transmission of the pressure wave 
varies inversely with the vascular capacitance of the arteries. Also, it 
is known that the velocity increases progressively as the pulse wave 
travels from the ascending aorta toward the peripheral regions. The 
arterial pressure signal contour becomes distorted as the wave is 
transmitted down the arterial system. There are three major changes which 
occur in the arterial pulse contour as the pressure wave moves forward. 
The first is that the high frequency components of the pulse, such as 
those corresponding to the dicrotic notch, are filtered and soon 
disappear. The second change is that the systolic portions of the pressure 
wave become narrowed and elevated. Third, a "hump" may become prominent on 
the diastolic portion of the pressure wave. 
In elderly patients with less compliant arteries, the pulse wave may be 
transmitted virtually unchanged from the aorta to the periphery. The 
reason for this change is controversial. A common explanation is based on 
the concept that the pulse wave is reflected from branch points back 
toward the aortic arch to thereby set up pressure oscillations. In an 
elastic tube, a traveling pressure wave is reflected to some extent 
wherever there is discontinuity in the system. If the tube is completely 
blocked, reflection of the pulse wave energy is complete (and will be 
180.degree. out-of-phase). If a pressure wave travels with increasing 
velocity toward the periphery and reflects back from regions where many 
branches occur over a short distance, the oncoming pressure wave is 
distorted, attaining a higher peak pressure and wider fluctuations 
following the peak, while at the root of the aorta, the initial upstroke 
of pressure is extremely rapid. 
During the remainder of the systolic period (or "systole"), the pressure 
wave is rounded or "dome" shaped. The end of systole is clearly marked by 
a sharp dicrotic notch accompanying closure of semilunar valves. During 
the diastolic run-off, the pressure declines almost linearly. The pressure 
descends rapidly, and during the run-off period there is an additional 
wave called the dicrotic wave. 
In addition to distortion of the waveform resulting from reflected waves, 
changes in the pulse waveform can be visualized in terms of its frequency 
contents. The transmission velocity of the high frequencies of the signal 
is faster than that of the low frequencies. Under these circumstances, the 
more rapidly traveling high frequency waves may produce increased peaking 
of the pressure pulse and corresponding deformation of the remainder of 
the pulse. 
Various methods are known in the art for measuring cardiac output, a common 
approach involving a thermodilution technique in which a solution colder 
than body temperature is injected into the right atrium through a catheter 
and the resulting drop in blood temperature at the catheter tip indicates 
an amount of blood flowing around the tip. This method is invasive, 
however, and thus involves risks to the patient as it has the possibility 
of damaging the anatomical structures through which the catheter is 
threaded. Complications associated with pulmonary artery catheterization 
include pulmonary artery rupture, balloon rupture, sepsis, air embolism, 
etc. 
Other conventional ways of determining cardiac output include thoracic 
bioimpedance and continuous wave Doppler ultrasonography. These two 
techniques are non-invasive, and are thus preferred for patients with 
high-risk vulnerability to the invasive procedures such as thermodilution 
estimation methods. In the bioimpedance method, changes in resistance to 
microcurrents injected into a patient are measured in order to calculate 
stroke volume, i.e., the amount of blood pumped in a single beat of the 
heart. The method therefore involves a pulse-by-pulse determination of 
cardiac output whereby four pairs of surface ECG electrodes are placed on 
the neck and chest of a patient. The outer pairs of electrodes inject a 70 
KHz, 2.5 mA current into the thoracic tissue and the current is then 
sensed with the inner pairs of electrodes. The resistance to the injected 
current is dependent upon the fluid characteristics of the thoracic 
volume. Pulsating changes in thoracic resistance (bioimpedance) are then 
timed to the ventricle electrical depolarization and mechanical systole. 
Continuous wave suprasternal Doppler ultrasound uses a Doppler transducer 
placed in the suprasternal notch directed toward the ascending aorta. The 
Doppler probe measures the aortic blood velocity. The integral of aortic 
systolic blood velocity multiplied by the cross sectional area of the 
aorta gives the stroke volume. 
Other methods of cardiac output measurement are based upon the Fick 
principle. This method is simply an application of the law of conservation 
of mass, whereby according to this principle, the rate of uptake or 
release of substance to or from blood at the lung is equal to the blood 
flow past the lung and the content difference of the substance at each 
side of the lung. This method is most commonly used with oxygen as the 
analyzed substance. By means of in-dwelling catheters, arterial and venous 
blood samples were obtained and these samples were analyzed on a blood gas 
analyzer to obtain the oxygen saturation and the partial pressure of 
oxygen. 
The Fick technique has the disadvantage that the measurements are complex 
and can often require an entire day of analysis before cardiac output can 
be ascertained. This makes the Fick technique undesirable in real-time 
applications where quick results are required in order to keep the patient 
in a stable condition. 
There has been disclosed in the prior art, a technique for monitoring 
system vascular resistance (SVR) using derivative calculations of a blood 
pressure signal of a patient. In this method, the signal is differentiated 
and the points of greatest slope in the systolic portions are determined 
and then divided into the pressure which exists at these points. Such a 
calculation is proportional to the system vascular resistance, and 
multiplication by a resistance factor will yield the SVR value. However, 
the present inventors have discovered that this method may lead to 
inaccurate results in that the analysis of the blood pressure waveform 
does not take into account various factors which may lead to undesirable 
results. The inventors have found that the arterial waveform inherently 
contains various "artifacts" which cause errors in measurement of cardiac 
output. These "artifacts" are due to variable characteristics of the 
cardiovascular system such as different arterial capacitances of patients' 
arteries, reflected waves which cause aberrations in the signal, and also 
various damping characteristics of the waveform. There has thus been a 
need in the prior art for a technique which overcomes these drawbacks. 
SUMMARY OF THE INVENTION 
It is therefore an object of the present invention to provide a new method 
and apparatus for examining and indicating the status of the 
cardiovascular system of a subject, cardiac output, SVR and all the 
hemodynamic data which can be calculated from the parameters associated 
with the arterial waveform such as stroke volume and the index data of 
cardiac output. 
Another object of the present invention is to provide a method and 
apparatus which can be used with non-invasive techniques for monitoring a 
blood pressure signal. 
A further object of the invention is to provide more accurate measurements 
of the cardiovascular parameters with a fast analysis time and an 
apparatus which is inexpensive to use. 
A still further object of the invention is to alleviate the need for having 
a trained operator to monitor cardiac output, and also to provide wide 
patient acceptability. 
The present invention accomplishes the above objects by providing a novel 
method and apparatus for analyzing a blood pressure signal in order to 
estimate the overall cardiac performance of a patient. A blood pressure 
monitoring transducer, which can be any of a number of known transducers 
which are commonly used in the art, continuously monitors the blood 
pressure signal of a patient which is then converted to a digital form. 
After low-pass filtering, a waveform analyzer extracts a number of 
different parameters associated with the particular blood pressure signal, 
and calculations are then performed based on these extracted parameters. 
The significance of the invention is that only the necessary parameters 
are extracted while the undesirable "artifacts" are ignored. The invention 
is therefore able to achieve more accurate estimations of cardiac output 
and SVR by analyzing the contours of the patient's blood pressure 
waveform.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
Referring now to the drawings, and more particularly to FIG. 1 thereof, 
there is shown a block diagram representative of the overall arrangement 
of the present invention. As shown in FIG. 1, a hardware connection of the 
various elements used in the calculation of cardiac output and SVR 
includes a transducer 20 which is part of a well-known blood pressure 
monitoring device. The transducer is connected to an A/D converter 30 
which is in turn connected to low-pass filter (LPF) 40. The output of LPF 
40 is connected to the input of waveform analyzer 50. Analyzer 50 is then 
connected to calculator 60 which can be, for example, a minicomputer. 
Calculator 60 also has a second input connected to input device 10, which 
can be a keyboard or any equivalent input device well known to those 
ordinarily skilled in the art. The output of calculator 60 is connected to 
display 70 for outputting the values calculated by calculator 60 in visual 
form. 
The overall scheme of the system of the inventive method and apparatus will 
next be described. The blood pressure monitoring device which includes 
transducer 20 is used for generating a signal waveform representative of 
the patient's pulsating blood pressure which is then input to 
analog-to-digital converter 30 and then low-pass filtered in LPF element 
40. The filter is used to smooth the electrical signal and remove noise 
components in the signal. The monitoring device can be either of the 
invasive or non-invasive type, and is not critical to the invention. 
However, as discussed above in regard to the invasive techniques for 
estimating cardiac output, the non-invasive techniques are often 
preferable as they involve less risk to the patient. It is to be 
understood, however, that the inventive method and apparatus can be 
practiced with either technique. 
Next, waveform analyzer 50 extracts various parameters of the arterial 
waveform which are used for performing calculations in calculator 60, as 
will be described in more detail below. The input device 10 enables the 
manual input of the body surface area (BSA) of a patient by the operator, 
while display device 70 is used for the output of the calculated 
parameters which are calculated in calculator 60. Body surface area is 
calculated in a known manner based on the patient's height and weight. 
Referring now to FIG. 2, there is shown an arterial waveform of the blood 
pressure signal as represented on a coordinate axis graph whereby the 
pressure in millimeters of mercury is located on the y-axis and time along 
the x-axis. This waveform can be displayed on an oscilloscope if so 
desired, but is not necessary to the invention. As can be seen from FIG. 
2, the waveform signal is a continuous chain of successive peaks (although 
only two are shown for purposes of simplification), and these peaks and 
valleys correspond to the systolic (i.e., dilation) and diastolic (i.e., 
contraction) portions of the patient's heartbeat. As can be seen from FIG. 
2, there are distortions which exist at the tail end of the diastolic 
pressure just before a pressure peak corresponding to the systolic portion 
of the waveform. Upon low-pass filtering, as shown in FIG. 3, such 
aberrations in the arterial waveform are removed. FIG. 2 shows two 
different peaks of a patient. Each peak has different values of systolic, 
diastolic and average blood pressure, as well as different stroke volume 
(SV) and SVR. Each signal is an independent unit on which all calculations 
can be performed. The outputs can be time-averaged by a user. 
FIG. 4 illustrates graphically the amplitude and rise integral parameters 
of the arterial waveform for use in the calculations performed in 
calculator 60. As shown, the rise integral is determined based on the area 
under the curve from the lowest point in the diastolic period to the peak 
pressure occurring at systole. The vertical distance corresponding to 
these two points is detected as the amplitude of the waveform. 
In FIG. 5, the signal "overshoot" integral is determined based upon the 
area under the waveform curve corresponding to the low point in the 
diastolic period up until the time at which the dicrotic notch is 
detected. "Overshoot" is a situation of pressure reflection from the blood 
arteries, which is dependent upon a specific artery and completely 
unrelated to systemic pressure. It should also be noted that arterial 
capacitance is a parameter related to arterial flexibility. The heart rate 
is shown as being equal to a set time period (i.e., 60 seconds) divided by 
the time interval existing between two successive peaks of the waveform 
signal. The signal integral is shown as being the area under the waveform 
curve from the beginning of the systolic period to the end of the 
diastolic period. 
FIG. 6 illustrates the signal overshoot time which is indicated by the 
period between the peak of the systolic portion of the signal to the time 
of appearance of the dicrotic notch. The signal rise time is defined as 
the period of time corresponding to the beginning of the systolic period 
to the peak pressure point. FIG. 7 illustrates first and second overshoot 
portions associated with the run-off period during the diastolic portion 
of the signal. FIG. 7 shows the overshoot (which is the actual systolic 
pressure of the signal), where the overshoot appears similar to an 
additional notch. When such appears, its location will be between the 
actual dicrotic notch (2) and the peak of the signal. FIG. 8 shows a 
series of dicrotic notch points following the pressure peaks for the 
situation where there is no overshoot observed. Three consecutive signals 
are illustrated in FIG. 8. In FIG. 8 the notch is shown and there is no 
overshoot. 
FIGS. 9 and 10 show the first and second differentiations of the arterial 
waveform, respectively. As can be seen from FIG. 9, the high points of the 
first derivative correspond to the initial portions of the systolic period 
where the rate of change of the increasing pressure is greatest. 
Similarly, at the low point of the diastolic portion of the signal (where 
the slope is close to zero) the first derivative will show a low point 
which is then followed by a peak at reference point zero and then a 
decrease in the differentiated value. FIG. 10 illustrates that in the 
second derivative there a number of peaks on either side of reference 
point zero following each pair of high and low points associated with the 
peak pressure and low point in pressure of the diastolic portions. 
The cardiac output and SVR of the patient are obtained as follows: 
##EQU1## 
where .alpha. is a constant, s represents the minimum peak of the second 
derivative, h is the rise time of the signal, m is equal to the signal 
square root integral, n equals rise square root integral, f equals heart 
rate per minute, d is the amplitude of the waveform, BSA represents the 
patient's body surface area, a equals systolic peak pressure, i equals 
time interval between the point of maximum pressure of the systolic 
portion and the point at which the dicrotic notch is detected, p equals 
the maximum value of the first derivative, r equals the first maximum 
value of the second derivative, and t equals the second maximum value of 
the second derivative. 
##EQU2## 
where .alpha. is a constant, s represents the minimum peak of the second 
derivative, h is the rise time of the signal, m is equal to the signal 
square root integral, n equals rise square root integral, f equals heart 
rate per minute, d is the amplitude of the waveform, BSA represents the 
patient's body surface area, a equals systolic peak pressure, i equals 
time interval between the point of maximum pressure of the systolic 
portion and the point at which the dicrotic notch is detected, p equals 
the maximum value of the first derivative, r equals the first maximum 
value of the second derivative, and t equals the second maximum value of 
the second derivative. 
In deriving the above calculations for cardiac output and SVR, the values 
are calculated for several systolic peaks (e.g. five or more) and the 
CO/SVR values are averaged. The cardiac output and SVR algorithm of the 
invention allow the use of a wide range of known non-invasive techniques 
for producing the waveforms of the patient's blood pressure, such as 
photoelectric and piezoelectric methods. The most accurate results are 
obtained from an invasive method, such as the use of catheters placed into 
various arteries, particularly a radial or femoral artery. Invasive 
pressure monitoring is now routinely performed at the patient's bedside 
which therefore allows continuous measurements of blood pressure and 
analysis of blood samples. Such a method is the most accurate because the 
data produces a minimum of artifacts. Furthermore, the use of blood 
pressure data such as systolic, diastolic and pulse pressure can be 
displayed on a beat-to-beat basis. Also, the use of piezoelectric or 
electrooptic sensors with the inventive method and apparatus is possible 
when invasive blood pressure procedures are undesirable. The advantage of 
these techniques is that they are completely non-invasive, but they are 
also less accurate. 
Obviously, numerous modifications and variations of the present invention 
are possible in light of the above teachings. It is therefore to be 
understood that, within the scope of the appended claims, the invention 
may be practiced otherwise than as specifically described herein.