Apparatus and method of bioelectrical impedance analysis of blood flow

Apparatus and methods are provided for monitoring cardiac output using bioelectrical impedance techniques in which one or more interior electrodes are placed in the trachea in the vicinity of the ascending aorta, and one or more exterior electrodes are disposed near the suprasternal notch, so that the resulting bioelectrical impedance measurements reflect voltage changes induced primarily by blood flow dynamics, rather than respiratory or non-cardiac related physiological effects. Apparatus and methods are also provided so that the measured cardiac output may be used to control administration of intravenous fluids to a patient or to optimize heart rate for those patients having pacemakers.

FIELD OF THE INVENTION 
The present invention relates generally to apparatus and methods for 
non-invasively measuring cardiac output and more particularly to apparatus 
and methods for measuring cardiac output using bioelectrical impedance 
analysis techniques. 
BACKGROUND OF THE INVENTION 
Knowledge of cardiac output is crucial in the care of the critically ill 
patient, as well as patients with chronic heart disease requiring 
monitoring of medication. For many years the standard of cardiac output 
measurement has been pulmonary artery catherization. Previously known 
catherization techniques, as described, for example, in U.S. Pat. Nos. 
3,915,155, 3,726,269 and 3,651,318, involve periodic injection into the 
patient's bloodstream of a bolus of heated saline, during which 
thermodilution measurements are performed to determine cardiac output. 
Such techniques cannot generally be used for continuous monitoring. 
Moreover, such catherization techniques pose significant risk to the 
patient, including malignant arrhythmias, pulmonary artery rupture, and in 
rare cases, death. 
Consequently, for many years work has been underway to develop less 
invasive apparatus and methods for monitoring cardiac output. For example, 
as an alternative to catherization methods, Doppler ultrasound techniques 
have been adapted to measure the velocity of blood flow. Provided that the 
diameter of a vessel, its flow profile, and the angle of the ultrasound 
beam relative to the vessel can be determined, Doppler ultrasound 
measurements of the ascending aorta, either externally from the 
suprasternal notch, or internally from within the trachea, can be used as 
a measure of cardiac output. 
U.S. Pat. No. 4,671,295 describes an example of such methods and apparatus, 
wherein an ultrasound transducer is mounted on the tip of an endotracheal 
tube so that Doppler measurements of blood flow from a point (pulse wave 
mode) or path (continuous wave mode) along the ultrasound beam can be 
measured. The method described in the patent requires multiple 
measurements within the blood vessel, a priori knowledge of the blood flow 
pattern and cross-sectional area of the vessel, and the relative 
angulation of the blood vessel. In addition, the measurement is highly 
dependent upon the exact placement of the transducer. These drawbacks have 
resulted in the slow adoption of Doppler ultrasound cardiac output 
techniques. 
A yet further technique which the prior art has sought to apply to the 
measurement of cardiac output is bioelectrical impedance analysis (BIA). 
BIA has recently gained wide use as a method of measuring body composition 
and physiological metrics. BIA involves measurement of the electrical 
impedance (electrical resistance plus reactance) of body tissues as a 
function of the voltage drop experienced by a low level alternating 
current electric current that is passed through the body tissues between 
multiple electrodes. 
Generally, BIA apparatus employ two excitation electrodes between which a 
low level current is conducted, and two sense electrodes disposed at 
intermediate locations for sensing tissue impedance. Current flows 
predominantly through materials with high conductivities, such as blood. 
Less current flows through muscle, which has an intermediate conductivity, 
while the conductivity of fat, air and bone is much lower than that of 
either blood or muscle. Since the resistance to flow is a function of the 
conductivity and cross-sectional area of the conducting volume, volumes 
having a larger cross-sectional area provide lower resistance. 
It is also known that the impedance of the conducting volume and the 
measured medium metrics (i.e., static parameters such as fat or water 
content, and dynamic metrics, such as blood flow) are dependent upon the 
placement of the electrodes and the conducting path between the 
electrodes. Thus, the greater the distance between the electrodes, the 
more likely that extraneous variables will effect the measurement. 
BIA methods generally correlate the measured voltage drop between the 
electrodes using relatively simple algorithms based on simplified models 
of body structure, for example, by assuming that the body is composed of 
simple cylindrical resistive volumes. Temporal cyclical variations in the 
body impedance are then assumed to be a result of physiological events 
such as blood flow and breathing. 
Measurements of the electrical impedance, and particularly, how the 
electrical impedance varies with time, can therefore provide a 
non-invasive indicator of those events. Various algorithms have been 
developed to isolate the specific physiological parameters, such as 
cardiac output from the measured bioelectrical impedance, as described, 
for example, in W. G. Kubicek, et al., "Development And Evaluation Of An 
Impedance Cardiac Output System," Aerospace Medicine, Vol. 37, pp. 
1208-1212 (1966), which is incorporated herein by reference. 
Despite the application of BIA methods for measuring cardiac output, no 
simple continuous BIA-based cardiac output measurement device has gained 
widespread acceptance. Many existing BIA devices use external or internal 
electrodes to measure bioelectrical impedance for large volumes, for 
example, the whole body or thoracic segments. Because the sense current 
diffuses throughout the entire volume, making use of any and all 
conductive paths, differences between individual patients, and even for 
the same patient over time, may inhibit standardizing the BIA metrics. 
Moreover, it is known that while BIA measurements of cardiac output provide 
good correlation for normal patients and those hemodynamically stable 
patients, there is poorer correlation for critically ill patients and 
patients in heart failure, as described, for example, in R. J. Detemeter 
et al., "The Use Of Noninvasive Bioelectric Impedance To Determine Cardiac 
Output: Factors Affecting Its Accuracy," Am. J. Noninvasive Cardiol., Vol. 
2, pp. 112-118 (1988), which is incorporated herein by reference. 
An example of an attempt to overcome the variabilities encountered when 
taking bioelectrical impedance measurements across large volumes is 
described, for example, in U.S. Pat. No. 4,870,578. That patent describes 
BIA apparatus for monitoring cardiac output by using external electrodes 
that measure the electrical resistance of a segment of the thorax and 
includes circuitry to account for large voltage changes due to 
respiratory-induced voltage changes. As acknowledged in that patent, the 
respiratory-induced voltage changes are typically much greater than the 
cardiac-induced voltage changes. 
Other devices that attempt to account for the affect of non-cardiac 
physiological events on bioelectrical impedance include arranging multiple 
electrodes on catheters for insertion into the esophagus to measure 
thoracic bioelectric impedance, as described, for example, in U.S. Pat. 
Nos. 4,852,580 and 4,836,214. Both patents describe multi-electrode arrays 
inserted into the esophagus to provide an impedance measurement reflecting 
blood flow in the descending aorta. Such devices are not believed to 
provide true isolation of cardiac-induced voltage changes from those 
induced by other physiological events. In addition, these systems may be 
unable to provide positive contact of the multiple electrodes with the 
wall of the esophagus. 
In view of the foregoing, it would be desirable to provide apparatus and 
methods for accurately, non-invasively and continuously measuring cardiac 
output using BIA techniques. 
It further would be desirable to provide apparatus and methods for 
measuring cardiac output in critically ill patients using BIA techniques 
that overcome the inaccuracies arising from measuring voltage changes 
across whole body or large volume thoracic segments. 
It also would be desirable to provide inexpensive apparatus and methods for 
measuring cardiac output using BIA techniques that overcome the drawbacks 
of previously known BIA cardiac output measurement devices and methods. 
It would further be desirable to provide methods and apparatus for 
continuously monitoring cardiac output so as to permit the measured 
cardiac output to be employed as a metric for controlling and maintaining 
other aspects of a patient's health. 
SUMMARY OF THE INVENTION 
In view of the foregoing, it is an object of this invention to provide 
apparatus and methods for accurately, non-invasively and continuously 
measuring cardiac output using BIA techniques. 
It is another object of this invention to provide apparatus and methods for 
measuring cardiac output in critically ill patients using BIA techniques 
that overcome the inaccuracies arising from measuring voltage changes 
across the whole body or large-volume thoracic segments. 
It is yet another object of the present invention to provide inexpensive 
apparatus and methods for measuring cardiac output using BIA techniques 
that overcome the drawbacks of previously known BIA cardiac output 
measurement devices and methods. 
It is still another object of this invention to provide methods and 
apparatus for continuously monitoring cardiac output that permit the 
measured cardiac output to be employed as a metric for controlling and 
maintaining other aspects of a patient's health. 
These and other objects of the invention are accomplished in accordance 
with the principles of the invention by providing BIA cardiac output 
monitoring apparatus that measure only those volumes necessary to acquire 
cardiac output information. Apparatus in accordance with the present 
invention includes one or more interior electrodes placed in the trachea 
in the vicinity of the ascending aorta, and one or more exterior 
electrodes disposed in the vicinity of the suprasternal notch. Sense 
current conducted between the interior and exterior electrodes flows 
primarily through high-conductivity blood, so that voltage changes are 
induced primarily by blood flow dynamics, rather than respiratory or 
non-cardiac related physiological effects. 
Methods in accordance with the present invention overcome the inaccuracies 
of the gross physiologic models employed in previously known BIA cardiac 
methods, by avoiding the simplified algorithms for the ventricular stroke 
volume based on whole thorax BIA measurements. In particular, the methods 
of the present invention avoid the inaccuracies in whole body or thoracic 
BIA measurements associated with ignoring the multiple, branched and 
complex paths of blood flow. 
In accordance with the present invention, the capability to obtain BIA 
measurements in the vicinity of the ascending aorta, which has no 
branches, and which therefore directly reflects the flow of blood through 
the ascending aorta, provides a simple and highly accurate metric for 
computing ventricular stroke volume. 
In yet further aspects of the present invention, the apparatus for 
monitoring a patient's cardiac output may be used to control 
administration of intravenous fluids to a patient or to optimize heart 
rate for those patients having pacemakers. 
Further features of the invention, its nature and various advantages will 
be more apparent from the accompanying drawings and the following detailed 
description of the preferred embodiments.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
The present invention relates generally to the use of BIA techniques for 
measuring cardiac output in critically ill and heart-diseased patients. 
The apparatus and methods of the present invention overcome drawbacks 
observed in previously known attempts to use whole body or large volume 
thoracic BIA measurements to measure cardiac output, by providing 
apparatus and methods that are not based upon the gross modelling of 
physiological events implicit in such previously known BIA measurement 
techniques. 
In the exemplary embodiments of the apparatus and methods of the present 
invention, one or more interior and one or more exterior electrodes are 
disposed in close relation to the ascending aorta, so that voltage-induced 
changes can be closely correlated to cardiac events, without significant 
effects due to non-cardiac physiologic events. 
It is known in the medical literature that BIA measurements of cardiac 
output in general show good correlation for normal patients and 
hemodynamically stable patients, but much poorer correlation for 
critically ill patients, and patients in heart failure, as discussed in 
the above-mentioned Detemeter paper. Applicant has discovered that the 
reason for this poorer correlation in the latter cases is that the 
theoretical basis underlying the use of whole body or large-volume 
thoracic measurements may be incorrect. 
Previously known techniques derive the equation for ventricular stroke 
volume (SV) from the assumption that a time-varying column of blood, in 
parallel with the other conducting material in the thorax, changes from 
zero to a volume equal to the stroke volume during the cardiac cycle. The 
column of blood is assumed to be the length between the electrodes used to 
obtain the BIA measurements, with effects on the BIA measurements due to 
respiration accounted for, for example, as discussed in the aforementioned 
U.S. Pat. No. 4,870,578. 
Referring to FIG. 1A, derivation of a typical previously known BIA 
algorithm is illustrated. Cardiac output is estimated from the 
bioelectrical impedance measurement I(t), where it is assumed that changes 
in the bioelectrical impedance coincidental with the heart electrical 
activity (as represented by an electrocardiograph output) are the result 
of blood flow F(t). A transfer function T(t) is then based upon empirical 
formulae derived from measurements taken on healthy, hemodynamically 
stable subjects. The bioelectrical impedance is then computed as: 
EQU I(t)=T(t)*F(t)+N(t) (1) 
where N(t) is noise. 
Applicant has determined, however, that the foregoing assumption regarding 
the column of blood ignores the branched, multiple and complex paths 
present in the arterial system. Moreover, the distribution of blood and 
fluids between different physiologic "compartments" in the idealized 
thoracic or whole body model and body regions are different in normal and 
critically ill patients. 
As illustrated in FIG. 1B, the thoracic approach to BIA measurement must 
account for transfer functions appropriate to each of the multiple blood 
flow paths through the volume: 
EQU I(t)=.epsilon.F.sub.i (t)*T.sub.i (t)*W.sub.i +N(t) (2) 
where W.sub.i are weights corresponding to a priori knowledge of the 
relative distribution of flow through the various segments of the volume, 
e.g., the aorta, and arterial segments and other fluid chambers. Moreover, 
the weights W.sub.i may be different for different patients, may be 
different for chronically ill as opposed to healthy subjects, and may be 
variable even within a given patient, e.g., due to changes in heart rate. 
Applicant has therefore discovered that equation (1) can be used accurately 
for any patient provided that the transfer function T(t) is correlated to 
measured blood flow (e.g., using a flow meter) where the effect of the 
distribution weights W.sub.i can be essentially eliminated. Accordingly, 
applicant has concluded that BIA measurements should be taken very close 
to a blood vessel, so that between the electrodes of the BIA apparatus 
there are no branching vessels or adjacent vessels. The present invention 
therefore involves the use of BIA measurements in the vicinity of blood 
vessels meeting the foregoing requirements. 
Referring to FIG. 2A, the upper portion of a human body 100 is shown in 
outline, with the corresponding location of the aorta 101, esophagus 102 
and trachea 103 shown in dotted line. These vessels and organs are more 
clearly depicted in FIG. 2B. With reference to FIGS. 2A and 2B, the 
outflow tract of the left ventricle of the heart is the ascending aorta 
101a. Segment 101b of the artery passes in front of trachea 103 and up 
towards the base of the neck, then arches 101c towards the lower part of 
the body. 
Applicant has determined that the location of ascending aorta 101a, just 
beneath the suprasternal notch 100a, provides ready access for an external 
BIA electrode. In addition, because ascending aorta 101a passes directly 
in front of trachea 103, it is possible to obtain a BIA measurement across 
ascending aorta 101a with relatively little intervening tissue. And, 
because the first branches from the aorta are from aortic arch 101b, the 
measurement of blood flow from ascending aorta 101a accurately measures 
the volume of blood ejected from the left ventricle. 
Referring now to FIGS. 3 and 4, the apparatus of the present invention is 
described. Apparatus 10 comprises an electrode-bearing endotracheal tube 
12, one or more external electrodes 14, impedance recorder 16, digital 
sampler 17 and computer 18. 
As illustrated in FIG. 3, endotracheal tube 12 may comprise a standard 
endotracheal tube 20 having one or more electrodes 22 disposed on its 
distal tip 24. Endotracheal tube 12 may also comprise eccentric balloon 
cuff 26 to ensure good electrical contact between electrode 22 and the 
interior wall of a patient's trachea. Balloon cuff 26 includes 
conventional inflation means in fluid communication with balloon cuff 
through a lumen within endotracheal tube 20. Where balloon cuff 26 is 
employed, endotracheal tube 20 also includes central lumen 25 to provide 
ventilation to the patient as well as to permit the administration of 
oxygen. 
Electrode 22 may comprise a 6 mm conductive foil strip, for example, Type 
M6001, available from the 3M Company, St. Paul, Minn., which is 
electrically coupled to impedance recorder 16 via electrical leads 28 
disposed within, or outside of, endotracheal tube 20. Electrode 22 may be 
attached directly to the exterior of the endotracheal tube 20, so that the 
eccentric nature of balloon cuff 26 urges electrode 22 against the 
tracheal wall. Alternatively, electrode 22 may be disposed on the exterior 
of balloon cuff 26. 
External electrode 14 is placed in the vicinity of the suprasternal notch 
(see FIG. 1A) for voltage pickup, and is electrically coupled to impedance 
recorder 16 via leads 29. Electrode 14 may comprise a spot EKG electrode, 
for example, the AMI 1750-001, manufactured by Medtronic-Andover Medical, 
Boston, Mass. Additional BIA measurements may be achieved using additional 
electrodes 14 spaced apart from one another in the vicinity of the 
suprasternal notch. 
Impedance recorder 16 may be a commercially available impedance recorder 
providing both the sense current (generally less than 1 mA at a frequency 
of 50-100 kHz) and impedance measuring capability, for example the 
Minnesota Impedance Cardiograph Model 304A operating at 100 kHz. Signals 
output from the impedance recorder are digitally sampled by digital 
sampler 17, for example, at a rate of 250 Hz using a standard 12-bit 
analog to digital converter, available from ComputerBoards, Inc., 
Mansfield, Mass. The sampled output of digital sampler 17 is then provided 
to computer 18, for example, an IBM-compatible personal computer having an 
Intel 386 or higher microprocessor, for storage and processing, as 
described hereinbelow. 
Referring now to FIG. 5, arrangement of the electrodes 22 and 14 are 
described. Endotracheal tube 20 is inserted into patient 100 through nasal 
cavity 104, past epiglottis 105 and into trachea 103 in accordance with 
standard medical practice. If the apparatus of the present invention is to 
be used for only a relatively short period of time, e.g., while a patient 
is anesthetized during surgery, endotracheal tube 20 may be inserted into 
the trachea via the mouth. Alternatively, access to trachea 103 may be had 
through a surgical opening at the suprasternal notch 100a by conventional 
tracheotomy. 
One or more electrodes 14 are placed on the exterior of the patient near 
the suprasternal notch, so that the primary path for the excitation 
current is through the ascending aorta to electrode 22, with relatively 
little intervening tissue in the current path as compared to previously 
known BIA measurement techniques. 
Referring now to FIG. 6, the first derivative of the measured impedance 
(dZ/dt) (curve I) is compared to a typical electrocardiograph waveform 
(curve II) for a normal patient, where the components of the waveform 
describing events within the cardiac cycle are labelled. Curve I includes 
an A-wave component, due to atrial activity at the beginning of the 
cardiac cycle, represented by a downward deflection in the curve. The 
I-wave component represents an upward deflection in curve I occurring 
during isometric contraction. The B-wave component corresponds to the 
start of blood flow out of the ventricles, while the C-wave component of 
curve I represents the major upward deflection during the cardiac cycle. 
The amplitude of this deflection measured from the zero point is used in 
the calculation of the ventricular stroke volume SV. The X and Y points of 
curve I reflect closure of the aorta and pulmonary valves, respectively. 
Point O corresponds to rapid filling of the ventricles. 
SV is calculated according to equation 3: 
EQU SV=rho(L/Z.sub.0).sup.2 (dZ/dt.sub.p)T (3) 
where: 
SV=ventricular stroke volume, ml 
rho=resistivity of blood (in normal patients, about 150 ohm-cm/s, and can 
be corrected for each patient as a function of hematocrit) 
L=distance between the electrodes, cm 
Z.sub.0 =mean impedance between the measurement electrodes, ohms 
dZ/dt.sub.p =peak value of the upward deflection in the first derivative of 
the impedance waveform (amplitude of C-wave) 
T=ventricular ejection time (computed as the period between the occurrence 
of the B-wave component and point X in curve I). 
The digitized first derivative of the measured impedance is analyzed to 
extract the B-wave and C-wave components and the X deflection point. The 
amplitude of the B-C-X portion of the curve I waveform, and the time 
between these segments are then employed to compute stroke volume using 
equation 3. The distance between the electrodes L may be computed based on 
external body measurements. 
In a preferred embodiment of the invention, SV is continuously computed for 
each data segment that is of good signal quality, i.e., where the 
amplitude of the derivative of the impedance signal is above a certain 
quality metric. The SV may be continuously updated on a display (not 
shown) associated with computer 18, and may consist of a running average 
of the current and a user-selectable number of preceding cardiac cycles. 
Cardiac output may then be computed as the product of the instantaneous 
average SV and the heart rate, and also displayed numerically. 
Applicant expects that the transtrachealsuprasternal BIA measurement 
technique in accordance with the present invention will not be 
significantly affected by motion artifacts or electrode placement. In 
addition, because leads 28 and 29 can be relatively short, it is expected 
that the apparatus of the present invention will be less susceptible to 
electrical interference. 
In addition, as indicated hereinabove, alternative embodiments of the 
apparatus of the present invention may include multiple electrodes 22 on 
the endotracheal tube 24 and multiple external electrodes 14 for measuring 
the bioelectrical impedance. Alternative embodiments may also include 
additional sensors to enable additional quantitative analysis. For 
example, diodes suitable for employing blood oximetry techniques based on 
near infrared light absorption may also disposed on the internal and 
external sensors to measure blood oxygen saturation levels. 
Referring to FIG. 7A, another embodiment of the present invention is 
described as controller for fluids administration. In FIG. 7A, cardiac 
output is measured by apparatus 30 having two transtracheal electrodes 32a 
and 32b and two external electrodes 34a and 34b disposed on patient 200. 
Apparatus 30 functions as described hereinabove with respect to the 
apparatus of FIG. 4, and is used to monitor hemodynamic status and as a 
metric to control the administration of fluids intravenously via lumen 42 
coupled to fluid supply system 40. Computer 45, which may be an 
IBM-compatible PC (and the same computer that computes cardiac output from 
the measured values of bioelectrical impedance), controls fluid supply 
system 40. 
Operation of the apparatus of FIG. 7A is as follows. After a one unit loss 
of blood, for example, it is known that cardiac output changes but that 
heart rate and blood pressure do not. Thus, decreased cardiac output can 
be used to monitor the amount of fluids to be given to a patient. The 
apparatus of FIG. 7A provides a closed-loop system wherein the amount of 
fluid injected into the patient is controlled by the cardiac output 
computed as described hereinabove. 
In FIG. 7A, a baseline cardiac output measurement is obtained and then a 
bolus of 50 cc of fluid is given while cardiac output is measured 
continuously. As long as the cardiac output increases, additional boluses 
of fluid are given periodically, e.g., every 15 minutes. This process may 
be repeated several times a day for a critically ill patient. 
Referring to FIG. 7B, an alternative embodiment of the present invention is 
described in which apparatus 30 and computer 45 of FIG. 7A are used to 
control a pacemaker. It is desirable to maximize cardiac output for the 
lowest possible heart rate, since the lower the heart rate, the lower the 
myocardial oxygen consumption. In the arrangement of FIG. 7B, computer 45 
controls the setting of pacemaker 50 as described hereinafter. 
A baseline cardiac output measurement is first obtained and then the heart 
rate is reduced by a predetermined amount, e.g., two beats/min, while the 
cardiac output is continuously monitored by apparatus 30. As long as the 
cardiac output increases or remains unchanged, the heart rate is 
periodically further lowered by the predetermined amount, for example, by 
2 beats/min every 15 minutes. The process of reducing heart rate while 
monitoring cardiac output is continued until either a minimum desired 
heart rate is obtained or the cardiac output measured by apparatus 30 
begins to decrease. If the cardiac output is determined to have decreased, 
the heart rate is returned to the preceding higher rate. 
While preferred illustrative embodiments of the present invention are 
described above, it will be obvious to one skilled in the art that various 
changes and modifications may be made therein without departing from the 
invention and it is intended in the appended claims to cover all such 
changes and modifications which fall within the true spirit and scope of 
the invention.