NON-INVASIVE CARDIAC FUNCTION MEASUREMENT

Abstract: A method and an apparatus therefor for monitoring cardiac function in an animal or human subject including the steps of: placing a first movement detecting transducer on the torso, said transducer overlying at least part of two diametrically opposed borders of the heart or great vessels; generating a signal indicative of the movement of the torso portion subtended by the transducer, said signal including a cardiac component comprising at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform; and assessing cardiac function by monitoring changes in said ventricular volume waveform or said aortic pressure pulse waveform.

Description: NON-INVASIVE CARDIAC FUNCTION MEASUREMENT INVENTOR: MARVIN A. SAC NER, M.D. BACKGROUND OF THE INVENTION Field of the Invention This invention pertains to non-invasive monitors, and more particularly to non-invasive monitoring of cardiac function. Prior Art Although the electrocardiogram (EKG) has been the primary non- invasive device for continuously monitoring activity of the heart in clinical medicine, it reflects solely electrical activation of cardiac muscle and provides no information on the mechanical characteristics of the cardiac pump. Consequently, the EKG may show normal or near normal waveforms in the presence of greatly impaired blood pumping capacity of the heart. Conversely, the EKG waveform may be abnormal despite normal or near normal pumping action. In terms of life support, adequate circulation of blood from the heart to the tissues, as reflected by the blood pumping capacity of the heart, is of paramount importance. Obviously, non-invasive techniques for monitoring the blood pumping capacity of the heart are preferred over invasive ones. Nevertheless, invasive cardiac monitoring techniques, because of their perceived greater accuracy and ability to provide continuous monitoring, continue to be employed in, for example, critically ill patients. Invasive techniques generally have as their basis a catheter, such as a Swan-Ganz catheter, placed such that its tip lays within the pulmonary artery. This provides continuous recording of pressures in the pulmonary artery, and in certain instances pressures in the right ventricle, right atrium and indirectly the left atrium (pulmonary capillary wedge pressure) . Injection of inert dye or cold saline from the catheter allows discrete measurements of cardiac output by dye dilution method or ther odilution, respectively. Alternatively, sampling blood for oxygen content in the pulmonary artery and a systemic artery together with measurement of oxygen consumption permits calculation of cardiac output by the Fic principle. However, insertion of a cardiac catheter into the body may be hazardous. Its use can lead to death, which occurs in 1% of cases, and morbidity, which occurs in 33% of cases, as a result of infection and/or damage to the heart valves, cardiac arrhythmias, and pulmonary thromboembolism. Errors of technique, measurement, judgment and interpretation are common. It has been estimated that one-half million Swan-Ganz catheters used in the United States in 1986 resulted in the death of as many as 1000 or more patients. Furthermore, cardiac catheters cannot be kept in place for more than a few days owing to hazards from infection. They are also costly and labor intensive since catheterized patients require intensive care units which cost two to five times more than standard semi-private beds. In addition, health care workers face the risk of AIDS acquired virus and hepatitis virus as a result o exposure to blood of the infected patient during catheter introduction and subsequent maintenance. Moreover, cardiac catheters do not directly provide measurement of change in ventricular volume. While such measurements can be indirectly obtained in conjunction with injection of radiopaque dye and roentgenographic imaging, this technique is time-consuming and costly, and dangerous hypotension and bradycardia may be induced by the dye. Further, the number of studies in a given patient is limited by the hazards of x-ray exposure and radiopaque dye injections. Angiographic techniques provide the most widely accepted means for measuring ventricular volumes. They allow calculation of the extent and velocity of wall shortening and of regional abnormalities of wall motion. When they are combined with measurement of pressure, both ventricular compliance and afterload (i.e., the forces acting within the wall that oppose shortening) can be determined. When the results are expressed in units corrected for muscle length or circumferences of the ventricle, comparisons can be made between individuals with widely differing heart sizes. Cineangiography provides a large number of sequent observations per unit of time, typically 30 to 60 frames second. Although contrast material can be injected into pulmonary artery and left atrium, the left ventricle is outli more clearly when dye is directly injected into the ventricu cavity. Therefore, the latter approach is used in most patien except in those with severe aortic regurgitation in whom contrast material may be injected into the aorta, with resultant reflux of contrast material outlining the l ventricular cavity. Injection of a contrast agent does not produce hemodyna changes (except for premature beats) until approximately the si beat after injection. The hyperosmolarity produced by the contr agent increases the blood volume, which begins to raise preload heart rate within 30 seconds of the injection, an effect that persist for as long as two hours. Therefore, this technique can be utilized for repetitive measurements within a short time spa Further, contrast agents also depress contractility directl though newer nonionic agents have been found useful for minimizi these adverse effects. In calculating ventricular volumes or dimensions fr angiogra s, it is essential to take into account and app appropriate correction factors for magnification as well distortion produced by nonparallel x-ray beams. In order to app these correction factors, care must be taken to determi accurately the tube-to-patient and tube-to-film distance Correction is best accomplished by filming a calibrated grid the position of the ventricle. Thus, angiographic methods do n have wide clinical application owing to their complexity, safe considerations, invasiveness, and side effects of the contra agents. The importance of measuring changes of ventricular volume w well expressed by Davila in a symposium on measurement of le ventricular volume. He pointed out that the description of t functional mechanics of the left ventricle requires measurement force, strain and velocity (rate of strain) . Pressure, a standa measurement in cardiac catheterization laboratories, critical ca units and operating rooms, is not necessarily dependent on sha (geometry) or size (volume) of the ventricle. However, force a strain must be expressed in relation to geometry and size of t fluid container. In the same symposium. Chapman et al described cineangiographic method for measuring ventricular volume. The workers also took into account the shortcomings of their meth and made the following observations: "The ideal system f following change in ventricular volume is obviously one which fully applicable to the free-living organism, which requires n injection of any sort, and which can be used repeatedly over lon periods of time without danger or discomfort to the subject. Suc a system, if it ever becomes available, can hardly be based o roentgenologic principles. But until some entirely differen principle emerges and is applied, the roentgenologic principle i indispensable." A further requirement for an ideal system woul be a minimum of physician or technician time for utilizing suc technology and interpreting the results. Because of the obvious advantage of non-invasive technique over invasive ones, a continuing search has been made for reliabl non-invasive methods of assessing cardiac performance. Suc methods are needed particularly in detecting serial changes i cardiac function and in evaluating both acute and chronic ef ect of interventions such as drug therapy and cardiac operations. Th five principal non-invasive methods for assessing cardia performance are: systolic time intervals, M-mode and two dimensional echocardiography, radionuclide angiography, gate computerized tomography (CT scanning) , and gated magnetic resonanc imaging (MRI) . All but the first of these are alternatives t angiography for measurement of ventricular volumes and/o dimensions and therefore permit the non-invasive estimation o ejection phase indices. Other than in patients with obstructio to left ventricular outflow, wall stress (afterload) can b estimated from a combination of systemic arterial pressure, ventricular radius, and wall thickness. All four non-invasiv imaging methods allow estimation of ventricular systolic an diastolic volumes; none, however, is satisfactory for continuous or near-continuous monitoring of critically ill patients. Systolic time intervals have been usually obtained with the combination of an external transducer on the carotid artery in the neck to display its pulsations, a microphone over the heart to record heart sounds, and the electrocardiogram. This technique has never enjoyed wide popularity because of both technical and physiologic reasons: (1) reliable, reproducible recordings are difficult to obtain, (2) prominent internal jugular venous pulsations in the horizontal body posture may be superimposed on the carotid artery pulsations rendering interpretation of the carotid arterial waveform difficult, (3) accurate recording of heart sounds may be difficult to obtain particularly in patients with obesity or emphysema, (4) systolic time intervals are sensitive to many pharmacologic and hemodynamic influences including changes in left ventricular preload and afterload which may introduce misleading values, (5) changes in duration of systolic time intervals can be influenced by patient posture and time of day when recordings are made, (6) carotid pulse contours to calculate systolic time intervals can be difficult to interpret in patients with aortic valve disease, and (7) presence of congestive heart failure can either normalize abnormal values o make normal values abnormal. Echocardiography involves ultrasonic imaging of ventricula wall motion to monitor cardiac function. With this technique, th dynamics of ventricular wall contraction and the interna dimensions of the cardiac chambers can be recorded. Th apparatuses used for echocardiography encompass a wide variety o increasingly sophisticated and computer-aided imaging and analysi systems. The transducer placements on the chest require the services of skilled technicians and incorrect placements lead to misleading information. Furthermore, these systems are quite expensive, not readily portable, require that the patient be studied in the left lateral decubitus posture, and are not intended for continuous monitoring of critically ill patients throughout the day or during exercise. In addition to the foregoing drawbacks, echocardiography has several inherent limitations. For example, all ultrasonic beams have a defined breadth and height comparable to the size of the crystal transducer face. Beyond its focal point, the beam's cross- sectional area enlarges in direct proportion to the distance fro the transducer face. Therefore, in M-mode (single transducer) echocardiography, two laterally separated structures may appear in direct anteroposterior relationship. Two-dimensional electrocardiographic techniques also produc distortions, which increase with increasing distance between th target and the central beam axis. In these instruments, axia resolution (1-2 mm) is superior to lateral resolution (4-5 mm) Because of the complex nature by which two-dimensional images ar generated, artifacts may appear as intracardiac masses to th casual observer. Further, delineation of the endocardium of th left ventricle in its entirety is achieved only 70 to 80% of th time. Also, respiratory interference limits the ability to obtai continuous beat to beat recordings, particularly during exercise. Attempts have also been made to determine left ventricula end-diastolic and end-systolic volumes from dimensions derived fro echocardiography. These have met with variable success, dependin on the patient population studied and whether M-mode or tw dimensional echo techniques were employed. M-mode dimensions ar used to calculate left ventricular volume through an applicatio of the angiographic concept of the left ventricle as an ellipsoid. However, M-mode echocardiography allows measurement of only on left ventricular dimension, the septalposterolateral dimension, which is viewed at the level of the chordae tendineae. Consequently, to calculate volume from this single dimension, th following assumptions are made: (1) the ventricle being examine does in fact approximate the geometry of an ellipsoid, both i diastole and systole; (2) the septal-posterolateral dimensi measured coincides with the minor axis of the ellipsoid; (3) t orthogonal minor axis is equal to the measured minor axis; and ( the major axis is twice the length of the minor axes. While go correlations between angiographic and echo left ventricular volum have been obtained, correlations are poor in patients who ha asynergetic ventricular wall motion, which occurs in patients wi coronary artery disease in whom damaged areas of the le ventricular wall do not move in phase with the normal areas. Als because ventricular volume curves as a function of time cannot derived without utilization of several assumptions a approximations, they are not usually reported. Two-dimensional echocardiographyoffersconsiderableadvanta for estimation of left ventricular volume because it allows dire measurement of all three hemiaxes on the ellipsoid model and als allows application of other volume formulations, such as Simpson' rule. Studies have shown that correlations betwee echocardiographic and angiographic volumes are substantiall improved when two-dimensional methods are used, and goo correlations have been obtained even in the presence of ventricula asynergy. The greatest disadvantage to quantitative two dimensional echocardiography is the inability to obtain technicall satisfactory images in all patients and the labor involved i analyzing the studies. This technique, as with the M-mode, doe not readily provide dynamic changes of ventricular volume over time. Echocardiography has also been employed to estimate the velocity of ventricular circumferential fiber shortening (Vcf) . This echo measurement is analogous to the derivative of change in ventricular volume during systole and serves as a measure of ventricular contractility. Its application in M-mode echocardiography assumes that the left ventricular internal dimension is measured at the midventricular level. The mean rate of shortening is determined by dividing the calculated circumference expression by the left ventricular ejection time (ET) , which may be measured from the concomitant carotid pulse tracing or from the time duration of echocardiographic aortic valve opening. Peak Vcf can be similarly derived by extrapolation from the maximum systolic slope of posterior and septal walls. Vcf is inaccurate in patients with asynergetic movement of the left ventricle as in patients with ischemic heart disease. Mean velocity of circumferential fiber shortening (V cf ) can be determined simply from measurements of end-diastolic and end- systolic dimensions by echocardiography, CT scanning, or MRI. Since the ventricle is approximately circular at its minor axis the circumference is equal to diameter (D) . Mean V cf (in circumference/sec) is therefore the difference between end- diastolic and end-systolic circumference (in cm) divided by th product of the duration of ejection (in sec) and the end-diastoli circumference. Values of V cf obtained by echocardiography compar closely with those determined from cineangiograms. Echocardiography has also been employed to estimate strok volume (SV) , which is the difference between end-diastolic volum and end-systolic volume. This technique suffers from the inheren lack of accuracy in volume estimations and, clinically, strok volume varies widely with different physiologic circumstances suc as body size, heart rate, posture and exercise. It is, therefore not as useful a measurement as contractility. Nevertheless provided that subjects with left ventricular asynergy are exclude from analysis, fair correlations have been reported between strok volume derived from M-mode echocardiographic and two dimensiona echo techniques on the one hand, and both thermodilution an angiographic stroke volume measurements on the other. Another non-invasive technique is the apex cardiogram whic is obtained by employing a transducer over the maximal cardia impulse on the anterior surface of the left hemithorax i combination with the electrocardiogram. This technique is o limited usefulness for several reasons. In particular, th recording of the apex cardiogram is strongly affected by th characteristics of the recording transducer and coupling of th transducer to the skin surface. In the absence of a palpabl cardiac impulse on the chest, which may occur in patients wit emphysema, the apex cardiogram cannot be obtained. Moreover, interpretation of the apex cardiogram waveform for hemodynami measurements is even more problematic than systolic time intervals. Another non-invasive device for monitoring cardiac functio in the kinetocardiograph. This device records localized chest wall movements with a transducer consisting of a small metal arm attached to a flat end piece which directly contacts the chest wall. Motion of the metal arm is transmitted to a bellows, connected to a piezoelectric or strain gauge transducer. The bellows and pickup are mounted from a crossbar over the bed, and the end piece can be placed perpendicular to any location on the chest. The amplified signal, denoted the kinetocardiogram (KCG), is obtained during breath holding at end-expiration. The KCG measures low frequency inward and outward chest movements, which range from 5 microns in the left axilla to 200 microns directly over the precordium. Kinetocardiography differs from apex cardiography in which outward movements are accentuated by an air displacement funnel transducer placed over the apex of the heart (a position where pulsations can be felt by the examiner) . For example, the KCG senses true displacements of the precordium because of its extern crossbar frame of reference, whereas the apex cardiogram sens relative rib cage interspace motion. Also, the KCG is sufficient sensitive so that records can be obtained from many points over t precordium and not just at the apex as with the apex cardiograp KCG recordings in humans were initially described in locatio where the precordial electrocardiographic electrode leads we conventionally positioned. In these locations, the KCG general depicts inward motion of the chest wall following the QRS wave the electrocardiogram followed by a large number of low frequen vibrations superimposed upon an upward, outward motion. T investigators who initially described the KCG attributed the che movements to a combination of the following factors: (1) movemen due to the cardiac impact against the chest wall, (2) changes the intrathoracic blood volume as the result of ejection or filli of the heart, (3) impact of blood in the great vessels against t chest wall and (4) positional and shape changes of the contracti and relaxing heart. Tracings of KCG over the anterior a posterior rib cage reveal: (1) a carotojugular type of pul tracing in the infraclavicular area (attributed by t investigators to a mixed arterial venous pulse transmitted from t subclavian or axillary blood vessels), (2) with the subject pron a waveform configuration similar posteriorly to the electrocardiographic electrode placement position, and (3) wi upright posture, a smaller amplitude, noisy opposite deflection signal at a posterior position corresponding to the anterior KCG signal. The investigators attributed these findings to a combination of the factors listed above. The KCG depicts precordial outward systolic bulges in approximately 66% of patients with known myocardial infarctions. The largest outward motion is found most often at the V 3 electrocardiographic electrode placement position. Outward precordial bulges occur during exercise in about 30% of patients who develop anginal pain. Although the KCG appears to provide useful information on the mechanical properties of heart muscle, it has never received widespread clinical acceptance. This is probably because of: (1) the unwieldy transducer to patient interface; (2) restriction of patient movement and need for breathholding during recording; (3) noisy, often uninterpretable signals; (4) requirement of a great deal of skill to interpret recordings from different locations on the rib cage; and (5) lack of quantitation of the KCG waveforms with respect to changes of ventricular volume events obtained from analysis of the recordings. Another non-invasive device for monitoring cardiac function is the cardiokymograph (CKG). This device, available from Cardiokinetics, Seattle, Washington, consists of a circular, fl capacitive plate mounted in a plastic ring strapped to the ches Tissue motion beneath the transducer distorts an induc electromagnetic field which in turn alters the frequency of t oscillator plate. This change of frequency is converted to change of voltage proportional to the chest wall motion at t transducer site and then displayed as an analog waveform. The C provides waveforms during breathholding quite similar in appearan to the kinetocardiogram. It depicts left ventricular wall moti abnormalities just like the KCG and therefore can be used improve the diagnostic accuracy of exercise testing as additional marker of yocardial ischemia. The cardiokymogram suf ers from the same limitations as t kinetocardiogram, namely, (1) an unwieldy transducer to patie interface; (2) restriction of patient movement and need f breathholding during recording; (3) noisy, often uninterpretab signals; (4) requirement of a great deal of skill to interpr recordings from different locations on the rib cage; and (5) la of quantitation of the CKG waveforms with respect to changes ventricular volume events obtained from analysis of the recording Electrokymography and radarkymography are still oth techniques for non-invasively monitoring cardiac function. T motions of the borders of the cardiovascular shadow obtained wi roentgen rays can be visualized directly on a fluoroscope by usin a photomultiplier tube to give a phasic analog signal from cycli variations in light produced by movement of the underlying hear border (electrokymography) , or from a video monitor of th fluoroscopic image and similar tracking technolog (radarkymography) . A graphic record of the segmental motion o the left heart border provides recordings which closely resembl the contour curve of changes in left ventricular volume over time Such technology can be utilized to diagnose localize segmental dysfunction of the ventricular wall. For example radarkymography has been used to diagnose ventricular wal abnormalities, including asynergistic and akinetic motion associated with acute myocardial infarction. Radarkymograph compares favorably with left ventricular cineangiography in th diagnosis of asynergistic myocardial contraction. However, radarkymography and electrokymography can be use only where an interface is visualized between the cardia silhouette and adjacent structures. Poor visualization i encountered in pulmonary fibrosis, pulmonary edema, pleura fibrosis and bony distortions of the rib cage. Dyspneic patient are difficult to study since extraneous motions of the hear caused by respiration introduce artifacts. Finally, both method subject the patient to exposure to Roentgen rays and this hazar prevents their use in situations requiring long term monitoring. A still further non-invasive technique for monitoring cardia function is impedance cardiography. It has long been recognize that the passage of a high frequency, low electrical current signa between electrodes placed on the heart or directed through th heart across the intact thorax produces changes of electrica impedance which varies directly with the length and inversely wit the cross-sectional area of the conductor. In impedance cardiography, detection of localized motion o the heart is highly dependent upon the placement of the electrodes To circumvent the problems of electrode placement, the entir thorax is treated as a conductor by placing exciting and receivin electrodes at the upper and lower borders of the thorax. Thi permits estimation of the magnitude of cardiac stroke volume as th difference in impedance between systole and diastole. Absolut values of cardiac stroke volume (amount of blood ejected by th heart per beat) are obtained by incorporating the rate of chang of impedance (an index of the velocity differences in pulse volume into an empirically derived equation. It is the derivativ waveform of torso impedance that forms the basis for it measurement by the commercial device, the Minnesota impedanc cardiograph, for calculating cardiac output. Although impedance cardiograms were initially recorded during breathholding to eliminate impedance changes superimposed by respiration, it has been found that ensemble-averaging of torso impedance waveforms using the R-wave of the electrocardiogram as a trigger pulse provides comparable waveforms during normal respiration in healthy subjects at rest and exercise and in critically ill patients. Because changes of transthoracic electrical impedance to detect changes of cardiac volume are highly dependent on electrode placement, segmental changes of cardiac volumes and accurate reproduction of volume contours over time cannot readily be recorded with such technology. On the other hand, treating all changes of hemodynamics of the entire thorax as a single conductor appears to provide reasonable estimates of stroke volume of the heart. It has also long been recognized that heart motion produces gas flow within the lungs, though the mechanism of this phenomenon has puzzled investigators for many years. One of the earliest researchers suggested that each heart contraction sent a volume of blood out of the thorax and the consequent negative pressure inside the affectivity rigid container caused an inflow at the mouth. Although this "aspirating" effect of the heart was subsequently well documented, the observation that the flow pulses were al present in open-chest animal preparations pointed to oth mechanisms. Cardiogenic flow pulses have been attributed to direct beati of the heart against the pulmonary parenchyma. Althou artifactually induced vascular pressure pulses produce fl oscillations in the airways, these oscillations can still be se in an airway of a lobe to which the lobar branch of the pulmona artery has been entirely obstructed. Furthermore, injection 25-50 ml of saline into the canine pericardial sac markedl diminishes all cardiogenic oscillations within intrapulmona conducting airways despite the presence of normal pulmona arterial pulsations. These observations suggest that neithe pulmonary vascular pulsations nor volume changes of the heart which should not be affected by a small pericardial effusion, we responsible for cardiogenic flow oscillations. The heart has an irregular shape and contracts with a twistin action; this results in a forceful thrust to some parts of th adjoining lung, whereas other parts follow the inward movement o the myocardium. It is these localized transient inflations an deflations which appear to produce intrapulmonary to-and-fro flo oscillations. Pericardial fluid tends to make the external surfac of the pericardial sac more spherical so that rotation or twistin of the heart no longer produces a thrust against the lung, thereb diminishing cardiogenic oscillations of the air columns. The actual redistribution of the flow pulse amon intrapulmonary airways originating from the heart depends upo relative impedance of the airways. Its magnitude depends upon th force and acceleration of the cardiac movement. However, apar from the heart movement, intrapulmonary factors must also influenc the pattern and extent of transmission of the pressure impulse an the zonal volume changes that it causes. Thus, whether a zon adjacent to the heart deflates or not, giving rise to a flow puls in the airways subtending it, depends upon its time constant. Th smaller its compliance and resistance, the more likely it is t respond to the cardiogenic pressure impulse by emptying. I contrast, if the time constant is high (e.g., due to increase airway resistance) , minimal emptying occurs during the time of th pressure cycle, resulting in smaller or absent flow pulses in th airways. The preceding discussion accounts for a number of experimental observations regarding recordings of expired gas flow. Thus, although cardiogenic oscillations appear on recordings of continuous expired gas concentrations in most normal subjects, patients with emphysema may not demonstrate this phenomenon. Absence of cardiogenic oscillations has been observed in patients with bronchial asthma, with oscillations reappearing after parti relief of the bronchial obstruction. Lung disease oscillations a not seen in the trachea unless they are also present within t lobar airways. . Luisada in 1942 reviewed the historical background for t designation, "pneumocardiogram n , and defined it as the recordi of pressure changes which occur in the air passages of the lung a consequence of the heart beat. He noted that graphic recordin of this phenomenon were published as early as 1861 in animals a in humans in 1876. He utilized a pressure sensing transducer fr one nostril while the subject breathed normally and employ electronic filtering to eliminate the slower respiratory wave He attributed the four positive and five negative deflections the resulting complex waveform to the following events: auricular contraction; 2) papillary muscle contraction; 3) fir ventricular wave; 4) peripheral pulse; 5) second ventricular wav 6) semilunar valve closure; 7) first diastolic wave; 8) tricusp valve opening; and 9) second diastolic wave. He believed that t multiple waveforms present in the pneumocardiogram were due to t difference between venous inflow to, and arterial outflow from, t thorax. Blair and Wedd in 1939 measured rib cage movements from a si below the sternum by recording pressure changes within a bello pneumograph manufactured by the Harvard Apparatus Company. The cardiogenic oscillations recorded during breathholding were attributed by the authors to excessive outflow of blood from the chest over inflow into the chest. They calculated this volume to be 30 ml by assuming that the recording below the sternum was representative of the entire thorax. Cardiogenic oscillations during breathholding have also been observed on analog signals from devices which display the total external movements of the respiratory system. Such oscillations were noted by Lee and Dubois in 1955 who enclosed a subject within an airtight chamber, the body plethysmograph. The subject first breathheld after inspiring air and small oscillations of pressure (calibrated as a volume) were sensed from the body plethysmograph with a sensitive pressure gauge. These oscillations were attributed to the heartbeat, but no significance was attached to the resulting complex waveforms by Lee and Dubois or by the present inventor. After the recording was obtained while breathholding on air, the subject inspired nitrous oxide (N 2 0) , a soluble gas, which was taken up by the pulmonary capillary blood flow. In 1961, Wasserman and Comroe modified the body plethysmographic technique of Lee and Dubois by substituting the subject's own thorax for the rigid body plethysmograph. Change in spirometric volume then reflected the exchange of gas molecules between alveoli and blood as long as thoracic volume remain constant. The latter was an important requirement of the metho Accordingly, to continuously monitor any movements of the chest abdomen which would invalidate this requirement, two mercury i rubber strain gauges were placed around the rib cage and upp abdomen and connected together to permit analog recording circumferential movements of the combined rib cage and abdomina compartments. Wasserman and Comroe believed that the cardiogeni oscillations observed with their method reflected changes i thoracic blood volume. They did not consider the oscillations t be related to changes in ventricular volume. The presen inventor accepted the interpretation given by Wasserman and Comro to the cardiogenic oscillations observed with their technique an used Wasserman and Comroe's results in a review paper o measurement of cardiac output by alveolar gas exchange. In 1965, Bosman and Lee utilized a body plethysmograph flowmeter method "to study the effects of cardiac contraction upo changes in lung gas volumes during breathholding both with th glottis open and closed." They reported and depicted curves wit multiple rises and falls from the body plethysmograph an pneu otachograph. They interpreted these complex waveforms a showing an excess of aortic outflow over venous inflow to th thorax during systole and a reverse during diastole. Using mor sophisticated technology, their work confirmed the findings o Blair and Wedd. SUMMARY OF THE INVENTION The present invention, which is sometimes referred to her as the thoracocardiograph or TCG, is based upon the discovery t during breathholding, small oscillations detected by sensors plac on the rib cage (RC) and abdominal (AB) surfaces and ordinari used to monitor breathing patterns closely resemble ventricul volume curves and aortic pressure pulses depending upon the respective placements on these surfaces. These sensors inclu those which measures changes of rib cage and abdominal dimension such as the respiratory inductive plethysmograph which measur changes in cross-sectional area; the inductance circumferenti transducer which measures partial cross-sectional area; the mercu in silastic strain gauge, bellows pneumograph, and differenti linear transformer which measure circumference and parti circumference; magnetometers which measure diameters; a partitioned pressure, volume and capacitance body plethysmograp which measure volumes. BRIEF DESCRIPTION OF THE DRAWINGS In the drawings: Fig. 1 is a diagrammatic representation showing the placement of wide band (left panel) and narrow band (right panel) transducers about the human torso; Fig. 2 is a graphic representation showing waveforms derived in accordance with the present invention in the supine (left panel) and standing (right panel) postures; Fig. 3 is a graphic representation similar to the left panel in Fig. 2; Fig. 4 is a recording from a semirecumbent normal subject using ensemble averaging to display an averaged vascular pulse and a ventricular volume curve with their corresponding derivatives; Fig. 5 is a graphic representation showing the use of curve fitting techniques to extract cardiogenic waveforms from raw data derived in accordance with the present invention; Fig. 6 compares waveforms derived using a narrow band senso and a single bellows pneumograph; Fig. 7 compares waveforms derived with a narrow band senso and a surface inductive plethysmograph; Fig. 8 shows waveforms derived, for different horizonta postures, using a wide band sensor at the lower rib cage an another wide band sensor at the mid-abdominal level; Fig. 9 illustrates waveforms derived using wide band sensor at the upper and lower rib cage placements and showing the effec of lung volume on ventricular volume curves; Fig. 10 shows the effect of the Valsalva maneuver o ventricular volume curves derived in accordance with the presen invention; Fig. 11 depicts waveforms showing the effect of exercise o stroke volumes; Fig. 12 depicts waveforms showing the effect of exercise o stroke volume with the subject in the supine posture; Fig. 13 shows the effect of amyl nitrite on ventricular volume curves derived in accordance with the present invention; Fig. 14 is a recording of a carotid arterials waveform and a left ventricular volume curve in a subject with ischemic heart disease; Fig. 15 is a waveform derived in accordance with the present invention and showing dyskinetic motion of a ventricle segment resulting from pulmonary hypertension; Fig. 16 is a graphic representation of a comparison of stroke volume measurements derived in accordance with the present invention and in accordance with the thermodilution method; Fig. 17 depicts waveforms derived using narrow band sensors in accordance with the present invention and showing ventricular volume curves derived from dogs with various band placements and body postures; Fig. 18 is a graphic representation comparing stroke volume as determined in accordance with the present invention and as derived using impedance cardiography; Fig. 19 is a graphic representation of cardiac output derived from impedance cardiography upon injection with terbutali and saline; Fig. 20 is a graphic representation for comparison with Fi 19 and showing cardiac output as derived in accordance with t present invention upon injection of terbutaline and saline; Fig. 21 is a series of recordings showing the effect externally pacing the right ventricle on the ventricular volu curve of a mechanically ventilated, anesthetized dog; and Fig. 22 compares waveforms measured during breathing of ro air (left panel) and an hypoxic mixture (right panel) . DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS I have discovered that during breathholding, cardiogeni oscillations derived from sensors placed on the rib cage an abdomen surfaces for displaying breathing movements differ i waveform configuration depending upon the location enclosed by th vertical height of the sensor. The invention will be particularl described with reference to the respiratory inductiv plethysmograph and its associated sensors, though as noted above, the present invention may be practiced with other devices used fo measuring dimensional changes at the rib cage and abdomen. The respiratory inductive plethysmograph is commercially available from Non-Invasive Monitoring Systems, Inc. (NIMS) under the trade names Respigraph and Respitrak and is described in United States Patent No. 4,308,872, the entire content of which is incorporated herein by reference. Basically, this apparatus comprises two coils of Teflon-insulated wire sewn onto elastic cloth bands encircling the rib cage and abdomen. The leads from the wires are connected to LC oscillator modules, or preferably a shared module, such that the inductance of the wires comprises the inductance element of the oscillator. Changes in the cross- sectional area of the rib cage and abdominal compartments result in changes in the inductance of the wires and hence changes in the oscillation frequency of the oscillator. The resulting signals fo the rib cage and abdominal compartments are demodulated an displayed as analog voltage signals. In respiration applications, these signals can be calibrated and summed to reflect absolut tidal volume. Fig. 1 shows placements of sensors employed with th respiratory inductive plethysmograph. The left-hand panel in Fig. 1 illustrates the placement of commercially available wide ban (WB) sensors, 10 cm in height, on the upper and lower rib cage (RC) and mid-abdomen (AB) . In the usual application of this device t non-invasively monitor breathing patterns, the sensor shown at the upper rib cage closely depicts the placement for respirator monitoring. For purposes of the present invention, the respiratory inductive plethysmograph was used with modified sensors. In particular, sensors as employed in the present invention were only 2.5 cm in height, such that each sensor subtended a narrower portion of the torso than the commercial wide band sensors shown in the left-hand panel in Fig. 1. The narrow band (NB) sensors used with the present invention are shown in the right-hand panel in Fig. 1. The xiphoid process of the sternum has been taken as the arbitrary point of reference for placement for the NB sensors employed with the present invention, as it is an easily recognized anatomic location which demarcates the caudal limit of the bon thoracic cage in the midline from the cranial limit of the sof tissues of the abdomen. While the invention will be describe herein in conjunction with the NB sensors, it will be apparent a this description progresses that sensors of any height may b employed, depending upon the information being sought. Fig. 2 shows, for a normal adult, waveforms traced fro polygraph recordings of the electrocardiogram (EKG) and the analo voltage signal from a narrow band sensor employed with respiratory inductive plethysmograph as taken during sequentia breathholds. The QRS complex of the EKG, labeled R in Fig. 2, marks electrical activation of the ventricles of the heart, whic precedes ventricular muscular contraction. As is well known, contraction of the ventricle causes ventricular volume to decrease as blood is ejected (systole) from the ventricles into the thoracic aorta and pulmonary artery. As indicated in Fig. 2, a single narrow band sensor was moved either above or below the xiphoid process at 2.5 cm intervals before each sequential breathhold. As also indicated in Fig. 2, the trial was repeated in both the supine and standing positions, left and right-hand panels in Fig. 2, respectively. The uppermost cranial border of the rib cage in a normal adult 187 cm in height, whose waveforms are depicted in Fig. 2, was situated +25 cm above the xiphoid process. The multiple tracings at the xiphoid proce denote repetitive tracings of sequential breathholds from polygra recordings at this site and demonstrate good reproducibility of t measurement. Of course, the tracings at the xiphoid process we taken with a narrow band sensor disposed between the "-2.5cm" a "+2.5cm" positions, i.e. over the xiphoid process, the xipho process tracings being shown between the "+22.5cm" and "+25cm" Fig. 2 simply as a matter of convenience. By making the recordings during breathholding, wavefo deflections due to respiration are eliminated. Accordingly, it known that the waveforms depicted in Fig. 2 are due physiological changes unrelated to breathing. Because the chang in rib cage and abdominal dimensions reflected by the signals sho in Fig. 2 are of considerably lower amplitude than those resulti from respiration, the gain of the respiratory inducti plethysmograph employed in generating the waveforms, t Respigraph, was adjusted to about ten to twenty times the ga setting for respiration applications. The narrow band sensors were placed such that recordin were ultimately obtained from almost all horizontal cross-section regions of the rib cage surface in the supine and standi postures. The configurations of the resulting waveforms depicted in Fig. 2, which I have found to be related to specif cardiovascular structures, distinctly differed depending upon th cross-sectional location subtended by the band. Thus, th cardiogenic signals from bands disposed at the level of the xiphoi process show a rapid decrease in volume (systole) following the wave of the EKG, which reached its nadir shortly before or afte termination of the T wave of the EKG, depending upon the precis location of the band. As shown, the diastolic phase of ventricula muscular relaxation is marked, at the xiphoid process, by a brie initial rapid increase in ventricular volume to a more gradual ris before reaching a peak plateau coincident with the next R wave This plateau continues slightly past the R wave before th downstroke of systole repeats itself. The configurations of the cardiogenic oscillations shown i Fig. 2 are, as shown, extremely dependent upon location of th sensor both in the supine and standing postures. The waveform taken at the xiphoid as depicted in Fig. 2 closely resemble th ventricular volume waveform as measured by sensors of length, diameter or volume surgically installed on the hearts of dogs, o from a cardiometer enclosing the isolated heart. Still referring to Fig. 2, the band placed +25 cm above th xiphoid, i.e. at the uppermost portion of the sternum, depicts a upgoing deflection following the R wave rather than a downgoin deflection as detected at the xiphoid process. It more closely resembles the waveform of the descending aortic pressure pulse a detected in the prior art using other techniques. From +2.5 t +17.5 cm above the xiphoid, the amplitudes of the signals in th supine posture diminish but still resemble ventricular volum curves. There are less marked variations of amplitude in th standing posture. For example, the amplitude of the wavefor recorded with a band placed +10 cm above the xiphoid in th standing position is approximately equivalent in amplitude to th waveform at the xiphoid process. As also seen in Fig. 2, th timing of the systolic downstroke following the R wave and it slope varies among the recordings taken at different location above the xiphoid. The waveforms of cardiogenic oscillations in the supin posture show an initial upward systolic deflection at the xiphoi location which is more pronounced -2.5 and -5 cm below the xiphoid This upward deflection denotes the period of isovolumetri contraction, a well documented phenomenon. At locations below th xiphoid, the mid-anterior sections of the band lie on the abdomina surface but the lateral and posterior sections overly the rib cage Therefore, changes in left ventricular volume are primaril recorded at these locations because the cardiac apex of the lef ventricular wall is located at the lowermost portion of the ri cage. Further, the slope of systolic ejection appears to b steeper at these locations below the xiphoid than above it. Thi is consistent with prior art observations that apical segment display a higher velocity of contraction than basilar segments. As is well known, during isovolumetric ventricular contractio immediately after electrical activation of the heart muscle, shortening of the long axis predominates such that the hear becomes more spherical and the transverse diameter toward the ape actually increases. This phenomenon accounts for the brief, ofte quite prominent, upward systolic deflections of isovolumetric contraction at the xiphoid, -2.5, and -5 cm band locations, and the diminution or absence of an upward deflection at this same point in time in the waveforms in locations from +2.5 to +17.5 cm above the xiphoid. This is consistent with the observation in canines that the isovolumetric contraction of the left ventricle varies in prominence depending upon the location where the dimensional gauges are surgically installed. The circumferential and length waveforms from the canine left ventricle as reported in the literature display prominent isovolumetric contraction which is strikingly similar to the human isovolumetric contraction waveform from bands placed from -2.5 to -10 cm below the xiphoid (See Figs. 1 and 2). The upward isovolumetric deflections are less marked in the standing posture presumably because greater longitudinal orientation of the heart due to gravity produces a lesser spherical cardiac shape at the onset of systole than in the supine posture. The timing sequence in Fig. 2 is consistent with fluoroscop imaging of the heart in which the observer perceives a wave muscular contraction from the cardiac base to apex. Similar timi of the initial changes in ventricular volume with systole has al been described with dimensions recorded during biplane corona cineangiogra s. Fig. 2 also shows that the amplitude of the chan in ventricular volume is less at the cranial than the caud portions of the rib cage. Since the base of the heart is locat more cranially than the apex, the finding of lesser changes volume is consistent with the conclusion that the band measures t horizontal sector of cardiac volume changes subtended by the heig of the band. So, if the atria and ventricles lie anatomically the same horizontal plane at a particular rib cage locatio summation of such signals would be expected. And, indee summation of the ventricular and atrial volume curves as report in the literature is consistent with the waveforms observed positions +12.5 to +17.5 cm above the xiphoid as shown in Fig. Thus, in these waveforms, the downstroke of systole is more gradu at the base of the heart than the apex because the atria are their diastolic period and rising in volume thereby cancelling part the ventricular systolic volume amplitude. Further, at t nadir of ventricular systole, the upward rounded curve represen the predominant peak of atrial diastole. Fig. 3 shows tracings from the bands 2.5 cm in height in th same subject whose waveforms are depicted in Fig. 2, but taken on week later at locations ranging from 15 cm below to 10 cm abov the xiphoid process. As seen from a comparison of Figs. 2 and 3 the appearance of the waveforms is consistent for recordings take at the same locations, but one week apart, evidencing goo reproducibility of the results. Referring to Fig. 3, the band placed -12.5 cm and -15 cm below the xiphoid on the abdomina surface show deflections more closely resembling the abdomina aortic pressure pulse. It should be noted that the -15 cm locatio was 2.5 cm above the umbilicus. Although the description thus far is based on waveform generated during breathholding, the display of the average waveform at any location can also be obtained during breathing b the well known technique of ensemble-averaging using the R wave o the EKG or the upstroke of a systemic arterial pulse obtained non invasively or invasively as a trigger to display solely th hemodynamic signals while eliminating the breathing waveform. Fig. 4 shows the ventricular volume curve together with th electrocardiograph and also the descending aortic pressure puls from the upper rib cage with the electrocardiogram using an averag of 50 heart beats. In FIG. 4, starting with the left panel, fro top to bottom, the first panel shows the carotid arterial waveform; the second panel shows the carotid arterial waveform derivative; the third panel shows the ventricular volume curve from TCG jus below the xiphoid process; and the fourth panel shows th derivative of TCG. On the right, from top to bottom, the firs panel shows the descending thoracic aortic pulse obtained from TC just above nipple level on the RC; the second panel shows it corresponding derivative; and the third and fourth panels show respectively, the left ventricular volume curve from TCG just belo the xiphoid process, and the corresponding derivative. The findin of an aortic pressure pulse on the recordings shown on the righ side of this Figure demonstrates heterogenicity of cardiogeni oscillations from different thoracic sites. The hatched lin depicts the EKG; the lowest panel displays the second derivativ of the EKG. The preceding description of varied waveform configuration of cardiogenic oscillations obtained with external sensors place on the rib cage and abdominal surfaces accounts for th inconsistencies and misinterpretations regarding previou recordings of these signals. Thus, the signal from a whole bod plethysmograph represents the sum of both positive and negativ deflections from the rib cage added to positive deflections fro the abdominal compartment. Similar mixing of signals is displaye on the sum signal from the rib cage and abdominal signals utilizin the respiratory inductive plethysmograph or bellows pneumograph i which transducers are placed upon both the rib cage and abdomina surfaces. And in a previous study using a single bellow pneumograph placed just below the sternum, the authors interprete the waveform assuming that this location was representative of th cardiogenic oscillations of the entire thorax rather tha reflecting cardiovascular events localized to their recording site. Fig.5 illustrates a further technique for obtaining waveform in accordance with the present invention during breathing. Referring to Fig. 5, the irregular waveform in the upper tracin shows the signal detected during breathing from a single narro band sensor connected to a respiratory inductive plethysmograph, with the band positioned at the xiphoid, which indicates that the band is positioned over the ventricles. This raw signal includes both a larger amplitude respiration component, and a smaller one due to cardiac function, the latter being the one of interest here. To remove the signal component resulting from respiration, the raw signal in the upper tracing of Fig. 5 is matched, using a conventional curve fitting equation with a cubic spline over sequential cycles, each of which comprises two cardiac beats. If this curve fit, depicted as the discontinuous "smooth" waveform in the upper tracing, is subtracted from the raw signal, the tracing depicted in the lower panel results, the discontinuities in the lower tracing resulting from the curve fitting technique described above, though these discontinuities may be eliminated by employing conventional smoothing techniques to adjacent curve fits as will be apparent to those of ordinary skill in the art. Similarly, "noise" on the lower tracing may be eliminated by high freque filtering. Even with the discontinuities and noise on the trac in the lower panel, it may be seen that the lower trac corresponds to ventricular volume curves as published in literature. The lowermost tracing in the lower panel is simply EKG. The removal of the respiratory waveform to provide beat beat display of the cardiogenic oscillations may also be carr out by other digital adaptive filtering techniques. Since compliance of the rib cage remains constant during br recording periods, change in amplitudes of the ventricular vol curve should provide accurate trending of relative changes stroke volume as well as ventricular contractility and relaxat characteristics. The product of stroke volume and heart r represents cardiac output and relative trends of the latter also available. Also, timing of systole and diastole slopes various portions of the ventricular volume curve, and vari volumes as a ratio to stroke volume, should allow comparisons am different subjects and trend plots over time in a single subje Finally, measurement of the absolute value of stroke volume independent methods such as dye dilution, thermal diluti impedance cardiography, radionucleide scans of the heart, echocardiography, angiography, etc. , allows one to set the init amplifier gains for the external sensor used in the pres invention to be equivalent to the values of stroke volume obtaine by the preceding methods. It has not been possible to calibrate the ventricular volum curve to absolute volumes independent of another method fo obtaining absolute values of stroke volume. However, it i possible to compare amplitude of cardiogenic oscillations from on site on the rib cage to another at a reference location. Thus, i a series of experiments involving six normal subjects, a band 2. cm in height was placed horizontally immediately below the xiphoi process and designated reference (REF) because solely the lef ventricle is transected anatomically at this site. Other band were placed 3 cm below REF, and 3 , 6, 9 and 12 cm above REF, an at the umbilical level. The electrical gain of respirator excursions of these bands was adjusted to be equivalent to the ban at REF and the amplitude of their cardiac waveforms was compare to the cardiac waveform of the REF band. In supine, semi-recumben and seated postures, at REF, 3 cm below and 3 cm above it, th cardiac waveforms had the contour of ventricular volume curves. More cephalad, waveforms tended to have complex oscillations. A the highest rib cage level and umbilicus, waveforms resemble descending aortic pressure pulses. Amplitudes of waveforms wer generally smaller at the +6 and +9 cm sites compared to the RE band in all postures, viz. 41% to 70% of REF (p < .01). There wa no correlation between amplitudes of cardiac and correspondin respiratory waveforms (r » -.14). Thus, this method of amplitu analysis should permit a study for obtaining normal values and capable of diagnosing hypokinetic ventricular segments (decreas motion) as might occur in patients with ischemic heart disease. As stated earlier, TCG appears to reflect changes in cros sectional area of the cardiovascular structures underlying t transducer. Since respiratory airflow and regional lung expansi may be altered by different density gases filling the lungs, investigated whether or not the TCG waveform was influenced by th factor. In addition to TCG for measurement of changes of stro volume (SV) , systolic and diastolic timing and volume event PEP/LVET was obtained as a carotid systolic time interval (STI) Six normal men breathed (1) air, (2) 20% 0 2 and 80% He, and (3) 2 0 2 and 80% SF 6 for 5 minutes and 3 measures of TCG and STI's we carried out over another 5 minutes. There were no differenc among the 3 gas mixtures whose densities varied 12-fold, in hea rate, SV, PEP/LVET, peak ejection rate/SV, and time of R-wave peak ejection rate. Therefore, this confirms that TCG measureme of ventricular function is unaffected by changes in physic composition of gases within the lungs. This is additional eviden that TCG displays changes in volume of underlying cardiovascul structures. While the present invention has thus far been described base on measurements taken with a respiratory inductive plethysmograp using narrow and wide bands, other externally placed respirator monitoring devices can be employed to record changes of cardia volumes and aortic pressure pulses. Fig. 6 illustrates this point The waveforms shown in Fig. 6 were obtained by placing narrow band connected to a respiratory inductive plethysmograph at the xiphoi and +25 cm above it, and a single bellows pneumograph (BP) +7. cm above the xiphoid process. NIP denotes the recording from neck inductive plethysmograph which provides a non-invasiv waveform of the carotid arterial pressure pulse as described i U.S. Patent Nos. 4,452,252 and 4,456,015, the entire contents o which are hereby incorporated herein. The EKG is also shown i Fig. 6. Fig. 6 shows that the waveform from the bellow pneumograph (BP) closely resembles the ventricular volume curv obtained using the respiratory inductive plethysmograph. Referring to Fig. 7, a recording taken with a surfac inductive plethysmograph (SIP) placed on the rib cage over th left border of the heart is shown together with the EKG and recording taken with the respiratory inductive plethysmograph a the xiphoid. As described in Canadian Patent No. 1,216,635, th entire content of which is hereby incorporated by reference, th SIP measures changes of surface cross-sectional area underneath th wire loop of the transducer. As seen in Fig. 7, the SIP als provides a recording depicting ventricular volume changes, thou the waveform appears slightly distorted compared to t corresponding waveforms obtained from cross-sectional slices arou the rib cage as recorded with the respiratory inducti plethysmograph using a band placed at the xiphoid process. Fig. 8 depicts, for different horizontal postures ventricular volume waveforms recorded with a wide band placed a the lower rib cage placement shown in the left panel in Fig. 1, an the abdominal aortic pressure pulse recorded with a wide ban placed at the mid-abdominal level shown in Fig. 1. The EKG is als shown in Fig. 8. The ventricular volume curves show simila configurations among the various postures, but there i accentuation of the isovolumetric contraction period of th systolic part of the ventricular volume curve in the left latera decubitus and the prone postures. Slight alterations i configuration with changes of posture are not unexpected since th heart is free to rotate and elongate in the rib cage as a functio of gravity. The sector of the heart subtended by the externall placed band or a similar external monitoring device would chang if the heart became oriented in a different plane. The abdomina aortic pressure pulse tracing is clearly recognizable in the supin posture and completely absent in the prone posture. This i probably because the supine posture permits maximum transmissio of the aortic pulse through the more compliant anterior abdomina wall whereas in the prone posture, aortic pressure puls transmissions to the anterior wall are highly damped leaving onl the back and the sides of the abdomen for transmission of vascula oscillations and the large amount of muscle mass present in th back and sides of the abdomen causes compliance (increase stiffness) of these regions which damps the aortic pulse pressur waveforms. Since compliance of the entire rib cage is much highe than the heavily muscled lower back, satisfactory recordings o ventricular volume curves are obtained in all horizontal postures Devices other than the respiratory inductive plethysmograp which are utilized to measure breathing patterns by changes o partial circumferences are conventionally placed on the anterio surface of the rib cage and abdomen compartments. These includ the bellows pneumograph, mercury in silastic strain gauges, an the linear differential transformer. They are incapable o accurately monitoring breathing movements in the prone positio because motion of the anterior surface of the transducer on th rib cage is restricted owing to the interposition of the transduce between tissues of the rib cage and the horizontal surface of th bed. Since these devices do not generally provide accurat measures of lateral and posterior motion, they cannot displa ventricular volume curves when the subject assumes the pron posture. Magnetometers, which are conventionally placed to measur changes of anteroposterior diameters of the rib cage and abdome compartments with respiration, do not produce accurat representations of changes of respiration nor ventricular volum when the subject assumes the lateral decubitus postures owing t exclusion of lateral rib cage movements which go undetected by th transducer. Fig. 9 depicts the effect of lung volume on ventricular volum curves obtained in accordance with the present invention. In eac of the four panels, recordings taken with a respiratory inductiv plethysmograph using a wide band at the upper and lower rib cag placements in the left-hand panel of Fig. 1 are shown together wit the EKG. In Fig. 9, near TLC indicates near total lung capacit signifying that the subject inspired a deep breath almost to th limit of vital capacity and breathheld with a closed airway at this lung volume level. FRC signifies functional residual capacity, i.e. lung volume at the end of normal expiration, and "between FR & TLC" in Fig. 9 signifies a moderately deep inspiration followe by a breathhold at this level. RV connotes residual volume, i.e. lung volume after full expiration, and "near RV" in Fig. 9 indicates breathholding at a lung volume near the lower limits of vital capacity. "Between FRC and RV" signifies breathholding afte a moderately deep expiration. As seen in Fig. 9, the configuration of the diastolic slop of the ventricular volume curve is altered by the lung volume level such that the terminal slope is flat at the high lung volumes and slopes upward at low lung volumes. Furthermore, the slope of initial ventricular systole is more gradual at the high lung volume ("near TLC" and "between FRC & TLC") than the steeper slopes at the low lung volumes ("near RV" and "between FRC & RV") . There are minimal differences in the amplitudes of ventricular volume curves at the various lung volume levels except for a slight increase at the level "between FRC & RV". These data suggest that myocardial contractility is increased during breathholding at low lung volumes compared to high lung volumes as expressed by the more rapid slope of systole at the low lung volume level. Furthermore, the flat slope of the terminal diastolic curve suggests that at the high lung volume levels ventricular compliance is decreased compared to ventricular compliance at the low lung volume levels. In the latter situation, the terminal curve slopes upward. This further suggests that the primary ventricular volume measured by the band at the lower rib cage placement is the left ventricular volume since it is known that both diminished myocardial contractility and lowered left ventricular compliance occur at increasing lung volume levels. Referring to Fig. 10, the effect of the Valsalva maneuver on ventricular volume curves derived in accordance with the present invention is shown. The Valsalva maneuver consists of straining against either a closed glottis or an occluded airway. Fig. 10 depicts such a maneuver with wide bands placed at the upper an lower rib cage placements depicted in the left-hand panel in Fig 1. Waveforms derived from the neck inductive plethysmograph (NIP for recording carotid arterial pressure pulses and the bellow pneumograph (BP) placed +7.5 cm above xiphoid for recordin ventricular volume are also displayed, as is the EKG. During th Valsalva maneuver, the pressure at the mouth rose to about 60 c H 2 0. The amplitudes of the ventricular volume waveform at th lower rib cage placement and the thoracic aortic pressure pulse a the upper rib cage placement showed a marked fall in amplitud during the Valsalva maneuver, as did the NIP and BP recordings The slope of systolic ejection of the ventricular volume curv during the Valsalva maneuver markedly diminished. The strok volume during the Valsalva maneuver for the band at the lower ri cage placement fell to 67% of the baseline, and rose 29% abov baseline upon release of the Valsalva maneuver. There was concomitant rise in the carotid arterial pressure pulse recorde with NIP, but the waveform of BP failed to disclose this rise. Th findings in Fig. 10 are similar to those obtained with lef ventricular angiography in which the fall in stroke volume derive from ventricular volume measurements fell from 35 to 75% of th baseline during the straining period. Similar declines of strok volume (53%) have also been obtained using an intracardia impedance catheter in the right ventricle. Echocardiograph measurements of ventricular volumes in both healthy subjects an patients with congestive heart failure demonstrate similar reductions in stroke volume during Valsalva maneuvers. The effect of exercise upon stroke volume depends upon the body posture in which exercise is carried out. In normal adults, utilization of dye dilution techniques for cardiac output allows calculation of stroke volume by dividing cardiac output by heart rate. During walking on a treadmill, one prior art study showed that stroke volume had an initial large rise with light exercise, i.e. heart rate rose from a baseline of 87 b/m to 115 b/m and stroke volume increased 69%. Stroke volume continued to rise slightly with more strenuous exercise up to a maximum of 84% above baseline at a heart rate of 171 b/m. On the other hand, in supine bicycle exercise, stroke volume increased only 6% during mild exercise, from a baseline heart rate of 71 b/m to 119 b/m. With moderate exercise, heart rate rose to 127 b/m but stroke volume increased only 13% over baseline. Referring to Fig. 11, in a normal adult instrumented with narrow bands connected to a respiratory inductive plethysmograph and seated on a bicycle, during breathholding immediately after terminating an exercise load, stroke volume increased from 35 to 65% over baseline on the band placed 2.5 cm above the xiphoid, while heart rate increased from 54 b/m up to 125 b/m. Referring to Fig. 12, in the supine posture, the rise of stroke volume with exercise measured with the band was much less than the seated posture, amounting to 32% over baseline, while heart rate increased from 67 b/m up to 116 b/ The increases of stroke volume exceed those reported for supi bicycle exercise using the dye dilution method for cardiac outp but are in agreement for the difference in response of stro volume to exercise in the seated and supine positions. As shown the rate of both systolic ejection and diastolic filling of th ventricle as measured with the bands markedly increased wit exercise. Amyl nitrite, a vaporized compound at room temperature whic is administered by nasal inhalation, produces an immediate fall i systemic vascular resistance associated with secondary alteration of cardiac hemodynamics. Fig. 13 shows the effect of this dru on the ventricular volume curve in a supine normal adult a reflected by measurements taken with a respiratory inductiv plethysmograph with a wide band place at the lower rib cag placement- and on the abdominal aortic pressure pulse as reflecte by measurements taken with a wide band at the mid-abdomina placement. Fifteen seconds after inhalation of amyl nitrite stroke volume increased 39% above baseline and heart rate rose fro the baseline of 54 b/m to 84 b/m. Myocardial contractilit markedly increased as indicated by the more rapid slope of systoli ejection after amyl nitrite. There was also a more rapid rise i filling during the diastolic portion of the ventricular volum curve. The increased rate of myocardial contractility was als present 27 seconds after amyl nitrite administration when the heart rate had slowed to 67 b/m. Thirty seconds after amyl nitrite administration, the heart rate was slower than baseline at 48 b/m and myocardial contractility returned to baseline value. Thus, measurement of ventricular volume curves with the respiratory inductive plethysmography in accordance with the invention provides a beat by beat recording during breathholding of the alterations expected from a drug which increases myocardial contractility and cardiac output. Of course, this information could also be derived during breathing by employing ensemble averaging or the curve fitting technique as more fully described above. Acute myocardial infarction with/without subsequent healing may produce paradoxical, dyskinetic or akinetic motion of the injured segment of the ventricle. In addition, silent ischemia may also induce such changes. Fig. 14 depicts a dyskinetic ventricular volume curve in a patient with ischemic heart disease. More particularly, FIG. 14 shows a recording of a carotid arterial waveform and left ventricular volume curve in a patient with ischemic heart disease BPs - systolic blood pressure; BPd = diastolic blood pressure; PEPu = pre-ejection period uncorrected for pulse waveform delay; LVET = left ventricular ejection time; dP/dt = maximum rate of rise of carotid arterial waveform. Ventricular wall dyskinesis is shown in the third recording from the top. Note that the time from the R wave of the EKG to Peak Ejection Rate (PER) is markedly prolonged to 520 s. Identica findings were obtained with echocardiography. Dyskinetic motio also may be present in patients with pulmonary hypertension in who right ventricular enlargement is present. This event has bee detected with bands located 5 cm above the reference band (place just below the xiphoid process) which showed a normal lef ventricular volume curve contour (Fig. 15) . Cardiac output can be measured by the thermal dilution metho using a Swan-Ganz cardiac catheter whose tip is placed within th pulmonary artery. Stroke volume is calculated by dividing th cardiac output by heart rate. This value can be used to calibrat the systolic portion of the ventricular volume curve (obtained wit a narrow band sensor or similar transducer) to an absolute volum value in the baseline period. Thereafter, this value can b utilized for all subsequent calculations of stroke volume from th ventricular volume curves to ascertain both the absolute volum variations and to establish the validity of the non-invasiv measurement. The accuracy of the latter depends upon th assumption that the ventricle can be considered as moving with on degree of freedom but this assumption can only be proven b comparing the thermodilution method (or other cardiac outpu method) to the measurements made with the non-invasive TC technique. This experiment was carried out in six anesthetize dogs. Baseline values were obtained by simultaneous collection o narrow band derived (TCG) and thermodilution values of stroke volume and cardiac output. The animals were then give 50 ml infusions of a 10% dextran 40 solution every 15 minutes with repeated simultaneous measurements at each time interval until cardiac output by the thermodilution method no longer increased with further dextran 40 infusions. Stroke volume by the thermodilution method rose to a maximum of 40% above the baseline value. In 46 paired comparisons, 87% of stroke volume values based on the TCG fell within 20% of stroke volume measurements based on thermodilution values (Fig. 16) . Therefore, TCG appears to provide an accurate measure of changes of stroke volume and cardiac output in anesthetized dogs. Because the rib cage of the dog is highly compliant much like the rib cage of a human baby, a study was undertaken to determine if it would be possible to obtain satisfactory recordings of ventricular volume curves using bands on the rib cage and abdomen in dogs. Fig. 17 depicts recordings of bands placed around the upper rib cage of the dog just underneath the axilla with variations of placement approximately 1 to 2 cm upward or downward. The left panels depict recordings when the dog was in the left lateral decubitus position on a flat table. The upper left panel shows a typical ventricular volume curve along with a superimposed rounded upward wave of atrial diastole at the nadir of ventricular systole. In the lower left panel, the band has been moved downward about 1 cm and the wave of atrial diastole is now eliminated. There is absence of the isovolumetric contraction phase of th ventricular volume curve in the upper left panel as indicated b the tracing labeled "RIP-upper RC (NB)", which connotes a narro band placed at the upper rib cage and connected to a respirator inductive plethysmograph, but a prominent upward deflection on th right upper panel. The bands on the mid-upper abdomen wer considerably dependent upon their placement site. When the dog wa placed on a V shape table to support the body in a differen orientation, the mid-abdominal band on the upper right panel showe a typical waveform of the abdominal aortic pressure pulse with a easily recognizable dicrotic notch. Other band placements also gave abdominal aortic pressure pulses. It is clear from the foregoing description that external monitoring with non-invasive sensors which measure rib cage and abdominal movements are capable of recording cardiovascular events in adults, babies and animals. With appropriate sensors and data processing, recording of segmental ventricular volume curves and aortic pressure pulses on a beat by beat basis during breathholdin is possible. Average waveforms can be obtained during breathing through the technique of ensemble-averaging, using as a trigger the QRS complex of the electrocardiogram or the upstroke of a systemic arterial pulse recorded either non-invasively or invasively. Alternatively, the curve fitting techniques and adaptive digital filtering techniques may be employed to extract the cardiogenic waveforms from the respiratory waveform. Further, waveforms of ventricular volumes and aortic pulses at different cardiac cycle lengths, and at various points in the lung volume level and airflow, can also be obtained. The technology described herein carries major implications in terms of physiologic, pharmacologic and clinical understanding of cardiac function and diagnosis of heart disease in adults, babies, and animals. It is impractical to measure stroke volumes and cardiac output with the invasive thermodilution technique in normal human subjects in order to establish the validity of the thoracocardiogram (TCG) for measuring changes of stroke volume (SV) and cardiac output (CO) . There is ample evidence in the literature that the impedance cardiograph can measure changes of stroke volume and cardiac output if body posture is fixed. This device uses an empirical equation and an assumption that the thorax can be treated as a single conductor to derive values of stroke volume. The waveform from this devices resembles an aortic pulse and is opposite in polarity to the ventricular volume waveform obtained with the thoracocardiograph. Values of stroke volume measured by the thoracocardiogram [SV(TCG)] were compared to SV measured by impedance cardiograph (IC) in six normal semirecumbent men after subcutaneous injection of 0.25 mg terbutaline to cause increased SV. On another day, 1 ml of saline was injected subcutaneously to serve as a control. Data from TCG and IC were collected every fiv minutes during a 30 minute baseline and 90 minutes after injection Maximum increases of SV and CO after terbutaline were 27% and 50% respectively; SV and CO were not altered by the saline injection In 288 paired comparisons, 91% of SV(TCG) values fell within 20 of SV(IC) (Fig. 18). Further, there was no statisticall significant differences between IC and TCG derived cardiac output at any time point in the preceding investigation (compare Figs. 1 and 20) . These data indicate that TCG derived ventricular volum curves accurately estimate changes of stroke volume and cardia output. The configuration of the ventricular volume curve provide indices of systolic and diastolic function of the heart. Fo systolic function, these intervals and volume ratios were compare to systolic time intervals which are well known timing measures o the carotid arterial pulse, to establish relationships for systoli function between these two different measurements of cardia contractility. In the first series of experiments, effects o terbutaline were investigated. Terbutaline has been purported to be a beta-2 adrenergi agonist but its administration is associated with a marked, sustained increase of cardiac output (CO) . The latter i attributed to systemic vasodilation and possibly enhance ventricular contractility. See Chest, volume 68, pages 616 et seσ.. 1975. To further characterize its action, several non- invasive cardiovascular monitoring techniques were employed. The left ventricular volume curve (LWC) was displayed as an averaged cardiogenic oscillation with the thoracocardiograph (TCG) . The respiratory signal was eliminated by an ensemble-averaging method. In addition to systolic and diastolic volumes from LWC (TCG) , other parameters were measured: (1) heart rate (HR) by EKG; (2) blood pressure (BP) by cuff auscultation; (3) systolic time intervals from the carotid pulse obtained by comingation of an inductive plethysmograph band around neck and the EKG; (4) dP/dt of carotid pulse after calibration with BP; and (5) ejection fraction (EF) by equation utilizing PEP/LVET (PEP = pre-ejection period and LVET = left ventricular ejection time) . In a 2 day crossover study, 5 normals received terbutaline .25 mg subcutaneously or saline and data were analyzed at baseline and peak response, 10-20 minutes after injection. Compared to saline, terbutaline produced significant rises over baseline in HR (20%) , LVETI (8%) [LVETI = left ventricular ejection time index], EF (16%) i stroke volume (28%) , cardiac output (54%) , peak ejection rate (PER) (61%) , dP/dt (70%) and left ventricular stroke work (27%) . Terbutaline significantly decreased diastolic BP (9%) , PEPI (20%) , PEP/LVET (31%) R to PER time (13%) [R = R wave and PER = peak ejection wave] and peripheral vascular resistance (43%) . Early diastolic filling flows, volumes and timing were not altered. The simplest and most consistent parameter of the systolic port of the LWC was shortening of R-PER time, easily recognized poi on the EKG and TCG waveforms, respectively. The systolic amplitude of the ventricular volume curve can utilized to estimate trends in stroke volume, and in conjuncti with an invasive technique such as thermal dilution or d dilution, or a non-invasive technique such as impedan cardiography or echocardiography, etc., the ventricular volu waveform can be calibrated to an absolute volume. Measuremen can be obtained with the standard (wide) or narrow band transduce of the respiratory inductive plethysmograph, but alternative other non-invasive sensing devices that have been utilized f measuring breathing movements can also capture cardiovascul events as a function of the height subtended by their transducer The respiratory inductive plethysmograph is preferable to oth such devices because it can provide accurate display of ventricul volume curves independent of postural changes in the horizont plane, whereas other respiratory monitoring devices general cannot accurately record ventricular volume curves in the prone lateral decubitus postures. The product of heart rate times stro volume equals cardiac output. With the invention, cardiac outp measurements can be obtained at rest and exercise in both norm and diseased subjects. These measurements can be recorded duri breathholding on a beat by beat basis or during breathing wi Page missing at the time of publication uppermost tracing depicts the unpaced ("UP") EKG and ventricul volume curve; here the heart rate was 112 beats/minute. ventricular pacing ("VP") heart rates below the unpaced heart ra (75 b/m, 92b/m and 107 b/m) , the ventricular volume waveform h a similar appearance and timing relation to the QRS as the unpac recording. However, at higher ventricular pacing rates (123 b/ 132 b/m and 184 b/m) , there is a marked delay relative to the Q due to paradoxic (dyskenesis) of the ventricular segment subtend by the inductive plethysmographic transducer. The ventricul volume curves of FIG. 21 were obtained with a band sensor plac at the xiphoid process of the dog. The recording of both the electrocardiogram and t ventricular volume waveform might help to differentia supraventricular tachycardia with aberrant electrical conducti from ventricular tachycardia in which the QRS complex indistinguishable. These two arrhythmias require different mod of management, as the ventricular tachycardia is an immediate lif threatening cardiac arrhythmiawhereas supraventricular tachycard with aberrant conduction is not. The differentiation may possible by three means: 1) recording of atrial diastole from band placed 10 to 15 cm above xiphoid (Figs. 2 and 3) in a mann analogous to jugular venous pulsations for recording of regul atrial contractions during supraventricular tachycardia, observing abnormal ventricular waveforms with timing and pha abnormalities along with dyskinetic wall motion as in ventricula pacing, or 3) loss of the isovolumetric contraction period of th ventricular volume curve. Although continuous electrocardiographic recording over 8 t 24 hours with a tape recorder (Holter monitoring) is often utilize to detect transient cardiac ischemia (impaired blood supply to th ventricular muscle) as reflected by alterations of the ST segmen and T wave of the electrocardiogram, it is well recognized that th usefulness of this approach is limited by artifacts and the majo applicability of Holter monitoring resides in diagnosis of cardia arrhythmias. However, segmental abnormalities of ventricular wal motion may precede electrocardiographic abnormalities. Thus, th combination of electrocardiographic Holter recording and segmenta ventricular volume waveforms with bands or other devices fo sensing rib cage movements should improve the diagnostic accurac of detecting myocardial ischemia both in patients who have ches pain and those who do not (silent ischemia) . Measurement of changes of cardiac output in patients who ar critically ill should help to guide therapeutic decisions, eithe through the use of appropriate doses of drugs and fluids, or wit surgical interventions. The utilization of ventricular volum waveforms to estimate stroke volume in critically ill patient provides information regarding the effects of intravenous flui challenge, i.e., if intravenous fluids are given and cardiac out increases, then the therapeutic intervention probably appropriate. On the other hand, if intravenous fluid administered and cardiac output remains the same or falls, then fluid challenge is probably inappropriate. This algorithm mi diminish the utilization of invasive Swan Ganz catheters placed the pulmonary artery which are also used to ascertain whether flu challenges are appropriate through cardiac output measurement a pulmonary arterial and left atrial (wedge) pressure recording The employment of such technology carries major risks to t patient, including death, and to the health care worker the ri of viral hepatitis and AIDS because of exposure to blood product Non-invasive monitoring in accordance with the present inventi poses no hazards to the patient nor health care workers while sti providing similar hemodynamic information. The ventricular volu waveforms along with the electrocardiogram can be obtained at t bedside or transmitted to a video-based central stati computerized display for data processing either through hard wi connections or telemetry. Although cardiac output is an important parameter in guidi management of patients, trending of systemic oxygen delivery (D may be a more valuable test. D0 2 is defined as the product cardiac output and arterial oxygen content. It signifies t volume of oxygenated blood delivered to the tissues. A fall in produced either by decreased cardiac output, decreased arteria oxygen content or both can cause tissue ischemia and tissue death Since arterial oxygen capacity is a function of the amount o hemoglobin in the blood, viz. 1.34 ml of oxygen can combine wit 1 gm of hemoglobin, one can calculate oxygen content by multiplyin the oxygen capacity of the blood by arterial oxygen saturation The latter can be obtained non-invasively by means of commercially available device, the pulse oximeter. If hemoglobi content of the blood is stable, then relative changes in D0 2 can b obtained by multiplying arterial oxygen saturation by cardia output. Thus, trends of D0 2 can be monitored non-invasively usin pulse oximetry and TCG. To illustrate the importance of D0 2 measurements, consider th effects of breathing a hypoxic mixture. It has been reported tha breathing a hypoxic mixture (F,0 2 = 0.1) for 7 to 20 minute increases heart rate (HR) 24%, stroke volume (SV) 16% and cardia output (CO) 38% compared to room air (6 publications, 64 normals) CO was measured by indicator dilution techniques. I extended suc observations by administering graded hypoxic mixtures for 1 minutes, viz. F,0 2 of .17, .15, .12 and .10 to 7 normals t establish dose-responsiveness for cardiac performance and oxyge delivery (D0 2 - Ca0 2 x CO) . SV and CO were measured with th thoracocardiograph (TCG) . In addition, oxygen saturation (Saθ 2 ) from pulse oximetry, ejection fraction (EF) from an equatio involving PEP/LVET, and minute ventilation (V,) from RIP we obtained. The table below lists mean Sa0 2 and fractional chang of other parameters compared to F,0 2 = .21 (Sa0 2 = 96%). In t table, an "*" denotes a statistically significant difference fr F I°2 " .21. Sao, SB SY CO EE R-PER 2, fi0 2 .17 90* 1.03 1.01 1.03 1.02 .97* 1.06 . .15 88* 1.05* 1.04 1.07 1. 05* .96* 1.16 . .12 76* 1.19* 1.10* 1.30* 1.08* .90* 1.22* 1. .10 67* 1.27* 1.19* 1.48* 1.08* .87* 1.20* 1. In the table, F,0 2 = fractional concentration of oxygen in g mixture (room air = .21); Sa0 2 = arterial oxygen saturation; HR heart rate; SV = stroke volume; CO = cardiac output; EF = ejecti fraction; R-PER = interval from R wave of EKG to peak ejection ra on TCG ventricular volume curve (Fig. 22) ; V, = minute ventilatio and D0 2 = systemic oxygen delivery. Changes of HR, SV and CO F,0 2 = 0.1 agree well with prior reported values. In normals, rose proportionally so that D0 2 was maintained constant with bri graded decrements of Sa0 2 . This illustrates the importance considering D0 2 rather than CO alone. Not surprisingly, there we no untoward symptoms in these subjects despite falls in Sa0 2 values as low as 55%. Estimation of D0 2 with decreased Saθ 2 normal and diseased states over prolonged time intervals nee investigation since D0 2 ultimately determines tissue viability. In contrast to grade hypoxia and terbutaline administratio experiments, head-up tilting of normal subjects produces decrease cardiac output and decreased cardiac contractility. The amplitud of the TCG derived ventricular volume curve may not accuratel reflect the fall in stroke volume owing to changes in the volume motion coefficient of the rib cage with major changes of bod posture as in changing from supine to upright postures. However the configuration of the curve is altered in an expected way an provides useful information on contractility, viz. instead of shortening of the R-PER interval as in hypoxia and afte terbutaline injection, head-up tilting causes the R-PER interval to lengthen, a finding consistent with decreased cardia contractility. The monitoring of trends in cardiac output during anesthesi using the non-invasive sensor placed upon the surface of the ri cage in patients undergoing peripheral or abdominal (i.e. non-ches related) surgical operations provides a valuable measure of cardia performance. It is well known that anesthetic agents and surgica interventions often deleteriously affect cardiac output. Evaluation of appropriate cardiac pacing rates and the effect of different pacing sequences on stroke volume is an importan consideration in cardiac pacemaker therapy. This can b accomplished by analysis of beat to beat stroke volume estimation from ventricular volume waveforms obtained with external senso placed on the rib cage. In addition, control of optimal paci rates through a servo loop can be accomplished by monitoring stro volume to reset the pacing rate for optimal stroke volu performance during exercise. This has already been carried out a research basis with an intracardiac placed catheter for beat beat changes of cardiac impedance. The monitoring of ventricular volume curves should also useful in evaluating changes of cardiac output in subjects confin to inaccessible environments such as the magnetic resonance imagi device, space capsules, diving bells, diving suits, high and l pressure chambers etc. Measurements of stroke volume during various mechanic ventilatory modalities should help to establish mechanic ventilator settings which least deleteriously affect cardi output. The ventricular volume waveform measured with extern sensors on the rib cage can be obtained during mechanic ventilation by the ensemble-averaging, curve fitting and oth adaptive digital filtering techniques as described above to extra the cardiac waveform. The Valsalva maneuver viz. straining with a closed glott decreases stroke volume as shown above from measurements of t ventricular waveform in a normal subject. The stroke volume normally increases after the straining maneuver is halted and the glottis is opened. Such a response may not occur in patients with heart disease and therefore the maneuver may help to differentiate normal subjects from patients with heart disease. In addition to using the respiration signal for monitoring breathing patterns in babies with near SIDS (Sudden Infant Death Syndrome) , monitoring of stroke volume and cardiac output from non- invasive determinations of ventricular volume curves as described above should aid in the early detection of cardiac abnormalities since it is known that bradycardia is often associated with apneas in these babies. Since the invention provides a mechanical indication of cardiac performance, it will be useful in establishing a timely diagnosis of death from cardiac standstill even though electrical activity of the heart may still be present. The rapidity of ventricular emptying as a measure of myocardial contractility can be obtained as the slope of the ventricular volume waveform from the external sensing device placed on the rib cage during systole or by taking an electrical analog or digital derivative of this waveform. The slope of rapid filling for the ventricular volume curve at the end of isovolumetric relaxation provides a measure of the mechanical characteristics o ventricular muscle. The slope of late diastole provides a measur as to whether the heart is filled, has limited diastolic reserve or has a great deal of diastolic reserve as indicated by a upwar sloping deflection of this portion of the curve. All th situations discussed in the preceding sections, regarding cardia output and stroke volume, apply for the importance of analyzing th configuration of the ventricular volume waveform to assess cardia performance. The configurations of the ventricular volume and aorti pressure pulses may be abnormal in patients with heart disease a rest, exercise, sleep, and with environmental stresses, e.g temperature, humidity, etc. The waveform of the ventricular volum curve in patients with valvular heart disease has distinctiv characteristics. For example, in patients with aortic stenosis th rate of systolic ejection of the ventricular volume curve i diminished whereas in patients with mitral stenosis the rate o diastolic filling is diminished. The upstroke of the aorti pressure pulse is also diminished in aortic stenosis. Patient with coronary artery disease may have limited ventricular wal motion due to ventricular compliance and have slow filling o diastole. Patients with constricted pericarditis or restrictiv myocardiopathy may show diastolic plateaus as a result of thes defects. A long flat diastolic plateau has been observed in t ventricular volume curve obtained with the present invention in patient with pulmonary edema, a pulmonary arterial wedge pressu of 27 mmHg, and an enlarged heart on the chest roentgenogram. Th type of waveform presumably indicates ventricular distention a might serve as a non-invasive monitor of left atrial pressure i such patients. Abnormal ventricular motion takes place with stunne myocardium after myocardial ischemia secondary to occlusion of coronary vessel or with therapeutic angioplasty in which brie occlusion of the coronary artery supplying a region of ventricula muscle produces abnormal wall motion of this part. Indeed abnormal wall function during myocardial ischemia precede electrocardiographic abnormalities and is a more sensitiv diagnostic sign. Acute myocardial infarction produces abnorma ventricular volume waveforms which may be reversed b administration of thrombolytic agents. This phenomenon is bes studied with segmental sensors over a large height of the rib cag rather than a wide band sensor enclosing the entire ventricle sinc small regions of abnormal motion might be missed under thes circumstances. The configuration of the ventricular volume curv during the Valsalva maneuver in which systolic ejection and strok volume are markedly diminished in normal subjects, and is followe by an increase of these parameters after release of straining, not occur in patients with heart disease and thus offers crite for distinguishing normals from patients with heart diseas Furthermore, changing the configuration of the ventricular vol curve by tilting the subject from the supine to upright postu and vice-versa produces characteristic alterations in t configuration of the ventricular volume waveform. For example, the standing posture, the terminal diastolic portion of t ventricular volume curve normally slopes upwards whereas in t supine posture terminal diastole has a flat plateau. Th signifies that the heart is well filled in the supine but not t upright posture, which might not occur in patients with hea disease. With narrow band external sensors, ventricular volu waveforms at different portions of the ventricle can be record such that timing and motion analysis between the segments can carried out. This should prove useful in assessing the effects acute ischemia and myocardial infarction on configuration of t ventricular volume waveform since it is well known that ventricul wall motion is impaired in these circumstances. This can resu in dyskinetic, akinetic or hypokinetic motion of segmental portio of the ventricular wall with consequent abnormalities of t segmental ventricular volume waveforms. Using the non-invasi method of the invention with sensors on the rib cage to displ segmental ventricular volume waveforms should make possible th diagnosis of such abnormalities and to ascertain the effectivenes of treatment either with intravenous administration of thrombolyti agents or angioplasty of the appropriate coronary artery. Furthermore, long term periodic follow-up with the non-invasiv technology of the invention should help in establishing th efficacy of treatment. For example, the effect of coronary arter bypass grafts on segmental ventricular volume curves can b determined post operatively; if follow-up evaluations show ne segments of abnormal wall motion different from the baselin established after surgery, then diagnosis of restenosis of th coronary artery might be suspected. Analysis of segmental ventricular volume configuration with such interventions as cardiac pacing, exercise, Valsalva maneuver, tilt, and drug administration, etc. should enhance its diagnostic effectiveness. The effects of anesthesia agents on ventricular volume waveforms should help to guide decisions on cardiovascular status during surgery. Finally, ambulatory Holter monitoring using the electrocardiograph and segmental ventricular waveform analysis with separation of curves into histograms of cardiac lengths and electrical abnormalities such as the ST-T wave depressions or inversions can be utilized to correlate electrical and mechanical events during arrhythmias and periods of potential myocardial ischemia. In conjunction with invasive catheterization of the le ventricle, ventricular pressure-volume curves can be construct to attain a definitive understanding of ventricular performance With an array of external transducers placed on the abdom of a pregnant woman and recording of the fetal electrocardiograp it should be possible to recognize and distinguish the sensor whi contains the waveform of ventricular volume by ensemble-averagi or adaptive digital filtering methods. The latter techniqu should eliminate maternal respiratory and cardiovascular pulsatio leaving only the ventricular volume curves of the fetus. Th measure would help to diagnose fetal cardiac distress by displ of both electrocardiographic and ventricular volume waveform musc abnormalities and provide early identification of fetal distre which might require obstetrical interventions. In conjunction with the ventricular volume curve, the analys of the thoracic aortic and the abdominal aortic pressure puls should provide useful information on diagnosis of valvular hea disease such as aortic stenosis and a convenient non-invasive mea to follow the outcome after surgical valvular repair. Thus, t upstroke of the aortic pressure curve will diminish with aort stenosis. Abnormal aortic pressure pulses occur with stable a dissecting aneurysms of the thoracic and abdominal aorta and shou help in establishing their diagnosis. In sum, the utilization of the non-invasive method of t invention for recording ventricular volume waveforms eith globally or segmentally together with analysis of aortic pressu pulses is an important advance in clinical and research cardiolog The electrocardiogram has served a highly useful purpose as indicator of normal and abnormal electrical activity of t heartbeat, but provides no information on the mechanical respons to electrical activation. The invention described herein is t first known to continuously non-invasively monitor mechanic performance of the heart by display of segmental characteristic It is also the first known invention to quantitatively continuous monitor changes in stroke volume. Further, the same extern transducer for cardiac monitoring can be utilized to no invasively, continuously monitor the breathing pattern. Sever of the many applications that such a safe, non-invasive diagnost tool will accomplish have been described above. Obviously, ma other applications will come to mind in the future, and according the above description should be construed as illustrative and n in a limiting sense, the scope of the invention being defined the following claims.

Claims: CLAIMS 1. A method for monitoring cardiac function in an animal or human subject comprising: placing a first movement detecting transducer on the torso, said transducer overlying at least part of two diametrically opposed borders of the heart or great vessels; generating a signal indicative of the movement of the torso portion subtended by the transducer, said signal including a cardiac component comprising at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform; and assessing cardiac function by monitoring changes in said ventricular volume waveform or said aortic pressure pulse waveform. 2. The method of claim 1, wherein said movement detecting transducer comprises a conductor disposed on said torso portion for movement therewith, movement of said torso portion resulting in corresponding changes in the self-inductance of said conductor. 3. The method of claim 2, wherein said conductor extends about said torso portion and subtends a finite height. 4. The method of claim 3, wherein said height is about 2.5 cm. 5. The method of claim 1, further comprising the step of removing the respiration component from said signal. 6. The method of claim 5, wherein said respiration component removing step comprises performing said assessing step during breathholding for avoiding changes in said signal due to respiration. 7. The method of claim 5, wherein said respiration component removing step comprises ensemble averaging said signal for removing the respiration component. 8. The method of claim 1, wherein said respiration component removing step comprises subtracting a curve fit from said signal for removing the respiration component. 9. The method of claim 1, wherein said respiration component removing step comprises adaptive digital filtering of said signal for removing the respiration component. 10. The method of claim 5, further comprising high pass filtering said signal for removing noise. 11. The method of claims 1 or 3, wherein said transducer is disposed at or near the xiphoid process and wherein said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform is said at least segmental ventricular volume waveform. 12. The method of claims 1 or 3, wherein said transducer is disposed at or near the uppermost portion of the sternum or the abdomen, and wherein said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform is said at least segmental aortic pressure pulse waveform. 13. The method of claims 1 or 3, wherein said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform is said at least segmental ventricular volume waveform, and wherein said assessing step further comprises monitoring the amplitude of said ventricular volume waveform for monitoring stroke volume. 14. The method of claims 1 or 3, wherein said assessing step further comprises monitoring changes in the slope, derivative of slope, or duration of said at least a segmental ventricular 15. The method of claim 13, and further comprising monitoring the heart rate of said subject; and monitoring changes in cardiac output by monitoring changes in the product of said heart rate and said stroke volume. 16. The method of claim 13, and further comprising measuring the absolute value of stroke volume by an independent method, and adjusting the level of said signal to indicate said absolute value, whereby said signal indicates absolute stroke volume. 17. The method of claim 1, further comprising placing at least one additional movement detecting transducer on the torso, said additional transducer also overlying at least part of two diametrically opposed borders of the heart or great vessels; generating a signal indicative of the movement of the torso portion subtended by said at least one additional transducer, said signal generated by said at least one additional transducer including a cardiac component comprising at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform; and wherein said cardiac function assessing step comprises monitoring changes in said waveforms generated by said first transducer and said at least one additional transducer. 18. The method of claim 17, wherein said assessing step further comprises comparing the waveforms of said signals generated by said transducers. 19. The method of claim 17, wherein one of said movement detection transducers is selected as a reference; extracting the respiration components of the signals generated by said transducers; adjusting the gain of the respiration component of each movement detection transducer other than said reference transducer to equal the gain of the respiration component of said reference transducer; and assessing cardiac function by comparing the amplitude of said cardiac component of said reference transducer to the amplitude of said cardiac component of at least one other transducer. 20. The method of claim 19, further comprising obtaining the relative amplitudes of said cardiac component of said signals generated by said transducers; repeating the steps of claim 19 on a known normal; obtaining the relative amplitudes of said cardiac component of said signal generated by said transducers when on said known normal; and comparing said relative amplitudes for said subject to said relative amplitudes for said known normal for assessing cardiac function of said subject relative to said normal. 21. The method of claim 1, wherein said movement detecting transducer is a bellows pneumograph, a mercury in silastic strain gauge, or a surface inductive plethysmograph. 22. The method of claim 1, further comprising generating an EKG signal for said subject, and wherein said assessing step further comprises monitoring changes in the timing of said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform relative to said EKG. 23. The method of claim 15, further comprising monitoring arterial oxygen saturation; and monitoring systemic oxygen delivery (D0 2 ) trends by monitoring trends in the product of cardiac output and arterial oxygen saturation. 24. The method of claim 1, further comprising measuring ventricular pressure; generating a signal indicative of ventricular pressure; wherein said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform is said at least segmental ventricular volume waveform; and constructing ventricular volume-ventricular pressure curves from said signals. 25. The method of claim 5, further comprising monitoring said respiration component. 26. The method of claim 1, wherein said subject is pregnant, and further comprising disposing said movement detecting transducer for detecting a cardiac component of said fetus; and removing the respiration and cardiac components of said subject from said signal. 27. The method of claim 26, further comprising removing said respiration component of said subject by monitoring during breathholding. 28. The method of claim 1, wherein said two diametrically opposed borders are the left and right borders of the heart. 29. The method of claims 17, 18, 19 or 20, wherein said transducers subtend an entire dimension of the heart. 30. The method of claim 29, wherein said dimension is the height of said heart from the most apical segment to the most basilar segment 31. An apparatus for monitoring cardiac function in an animal or human subject comprising: a first movement detecting transducer disposed on the torso, said transducer overlying at least part of two diametrically opposed borders of the heart or great vessels; means for generating a signal indicative of the movement of the torso portions subtended by the transducer, said signal including a cardiac component comprising at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform; and means for assessing cardiac function by monitoring changes in said ventricular volume waveform or said aortic pressure pulse waveform. 32. The apparatus of claim 31, wherein said movement detecting transducer comprises a conductor disposed on said torso portion for movement therewith, with movement of said torso portion resulting in corresponding changes in the self-inductance of said conductor. 33. The apparatus of claim 32, wherein said conductor extends about said torso portion and subtends a finite height. 34. The apparatus of claim 33, wherein said height is about 2.5 cm. 35. The apparatus of claim 31, further comprising means for removing the respiration component from said signal. 36. The apparatus of claim 35, wherein said means for removing the respiration component from said signal comprises means for ensemble averaging said signal. 37. The apparatus of claim 35, wherein said means for removing the respiration component from said signal comprises means for subtracting a curve fit from said signal. 38. The apparatus of claim 35, wherein said means for removing the respiration component from said signal comprises means for adaptive digital filtering said signal. 39. The apparatus of claim 35, further comprising means for high pass filtering said signal for removing noise. 40. The apparatus of claims 31 or 33, wherein said transducer is disposed at or near the xiphoid process and wherein said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform is said at least segmental ventricular volume waveform. 41. The apparatus of claims 31 or 33, wherein said sternum or the abdomen, and wherein said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform is said at least segmental aortic pressure pulse waveform. 42. The apparatus of claims 31, or 33, wherein said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform is said at least segmental ventricular volume waveform, and wherein said assessing means further comprises means for monitoring the amplitude of said ventricular volume curve for monitoring stroke volume. 43. The apparatus of claims or 31 or 33, wherein said assessing means further comprises means for monitoring changes in the slope, derivative of slope, or duration of said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform. 44. The apparatus of claim 42, and further comprising means for monitoring the heart rate of said subject; and means for multiplying said heart rate by stroke volume for monitoring cardiac output. 45. The apparatus of claim 42, and further comprising independent means for measuring the absolute value of stroke volume; and means for adjusting the level of said signal to indicate said absolute value, whereby said signal indicates absolute stroke volume. 46. The apparatus of claim 31, further comprising at least one additional movement detecting transducer disposed on the torso, said transducer overlying at least part of two diametrically opposed borders of the heart or great vessels; means for generating a signal indicative of the movement of the torso portion subtended by said at least one additional transducer, said signal generated by said at least one additional transducer including a cardiac component comprising at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform; and wherein said means for assessing cardiac function comprises means for monitoring changes in said waveforms generated by said first transducer and said at least one additional transducer. 47. The apparatus of claim 46, wherein said assessing means further comprises means for comparing the waveforms of said signals generated by said transducers. 48. The apparatus of claim 46, wherein one of said movement detection transducers is selected as a reference; further comprising means for extracting the respiration components of the signals generated by said transducers; means for adjusting the gain of the respiration component of each movement detection transducer respiration component of said reference transducer; and means for assessing cardiac function by comparing the amplitude of said cardiac component of said reference transducer to the amplitude of said cardiac component of at least one other transducer. 49. The apparatus of claim 48, further comprising means for obtaining the relative amplitudes of said cardiac component of said signals generated by said transducers, whereby cardiac function may be assessed by comparing said relative amplitudes for said subject to relative amplitudes obtained with a known normal. 50. The apparatus of claim 31, wherein said movement detecting transducer is a bellows pneumograph, a mercury in silastic strain gauge, or a surface inductive plethysmograph. 51. The apparatus of claim 31, further comprising means for generating an EKG signal for said subject, and wherein said assessing means further comprises means for monitoring changes in the timing of said at least a segmental ventricular volume waveform or a segmental aortic pressure pulse waveform relative to said EKG. 52. The apparatus of claim 44, further comprising means for monitoring arterial oxygen saturation; and means for monitoring systemic oxygen delivery (D0 2 ) trends by monitoring trends in the product of cardiac output and arterial oxygen saturation. 53. The apparatus of claim 31, further comprising means for measuring ventricular pressure; means for generating a signal indicative of ventricular pressure; wherein said at least segmental ventricular volume waveform or a segmental aortic pressure pulse waveform is said at least segmental ventricular volume waveform; and means for constructing ventricular volume-ventricular pressure curves from said signals. 54. The apparatus of claim 35, further comprising means for monitoring said respiration component. 55. The apparatus of claim 31, wherein said subject is pregnant, wherein said movement detecting transducer is disposed for detecting a cardiac component of said fetus, and further comprising means for removing the respiration and cardiac components of said subject from said signal. 56. The apparatus of claim 31, further comprising means worn by said subject for recording the transducer signal for accommodating ambulatory monitoring. 57. The apparatus of claim 46, further comprising means worn by said subject for recording said transducer signals for accommodating ambulatory monitoring. 58. The apparatus of claim 31, wherein said two diametrically opposed borders are the left and right borders of the heart. 59. The apparatus of claims 45, 46, 47 or 48, wherein said transducers subtend an entire dimension of the heart. 60. The apparatus of claim 59, wherein said dimension is the height of said heart from the most apical segment to the most basilar segment.

Inventors: SACKNER MARVIN A, SACKNER, MARVIN, A.
Applicants: NIMS INC, NIMS, INC.
Application year: 1990
Main classification: A61B5
IPC codes: A61B5