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Physiological correlates of impedance plethysmographic waveform.
Correspondence Address:
A review of investigations into the origin of impedance plethysmographic waveform is presented in this article. Attempts made by several investigators in the comparison of impedance plethysmographic estimations of peripheral blood flow with that obtained by standard methods are briefly described. Investigations indicating the negligible contribution from contact impedance at body electrode interface are highlighted. Temporal correlation of impedance plethysmographic waveform recorded from thorax with various important events of the cardiac cycle is summarised and various hypotheses on the genesis of this waveform are presented.
Peripheral impedance plethysmograms (IPG) using Nyboer's Technique[28] have been used for the estimation of peripheral blood flow for the past 5 decades. Though blood flow estimations obtained by either Nyboer's method or venous occlusion principle have correlated well with other standard methods, some of the investigators feel that the measured impedance changes are the result of artefacts produced by changes in contact impedance at the body electrode interface[14]. Some workers agree that the recorded impedance changes are due to altered electrical resistivity of the blood as the blood cells are not spherically symmetrical and they align themselves in the direction of flow[10]. Yet majority of researchers maintain that variation in the blood volume in a given body segment caused by blood flow is responsible to produce change in the impedance[28],[43]. Similarly, various hypotheses accounting for thoracic impedance change during a cardiac cycle are: (a) instantaneous changes in size, shape and movement of the heart, (b) blood flow pattern in vena-cava and the pulmonary veins, (c) cyclic perfusion of the pulmonary vascular bed and the pulsating pulmonary arterial blood flow, (d) pulsatile dilations of the aorta, (e) blood velocity related signals and (f) complex combination of simultaneous cardiac events[25]. Uncertain genesis of impedance change has hampered the wide clinical acceptance of impedance plethysmography., In contrast, volume displacement plethysmography, though less sensitive has been accepted as a routine clinical investigation, due to its direct relation with the volume of the body segment. Particularly in view of the gross assumptions made in the derivation of blood flow calculations, higher degree of experimental verification is needed for clinical acceptance of impedance plethysmography. In this paper, the authors review the significant contribution made by several investigators in this direction.
Jan Nyboer[28] resorted to in vitro study of forward flow, from a constant delivery pulsatile pump with a delivery valve, into a single expansile tubular segment. The fluid was collected in an open graduated cylinder through past flow hindrances of the outflow tube and the stop-cocks. He simulated the condition of arterial blood flow in this arrangement and obtained a correlation of 0.961 between actual flow and that estimated from IPG when there was minimum hindrance to outflow. Under the conditions of moderate hindrances, he obtained a correlation of 0.912. In his in vivo studies, Nyboer concluded that it was within the scope of impedance plethysmography to define the degree of hindrances to the blood flow in the extremities[28]. Van De Water et al[39] used a 5 lb liverwurst sausage as a model of limb, a section of dog's intestine being placed inside the sausage along its axis to simulate an artery. A perfusion pump was employed to pump saline through the artery. The four circumferential electrodes of an impedance plethysmograph system were applied around the sausage to measure the flow. For comparison the flow was also measured using an electromagnetic flow meter. They demonstrated a correlation coefficient of 0.942 between the flow estimated by IPG and that obtained by electromagnetic flow meter. Young et al[43] showed that the blood flow estimated by IPG using venous occlusion principle was 1.2 percent higher than that in the femoral artery in 5 dogs obtained by electromagnetic flow meter. By modelling the hind limb as a truncated cone and modifying the expression as A Z/Zo = 1.6 ?V/Vo to calculate the blood flow, they obtained a correlation of 0.980 between the flow estimated by IPG and that obtained by electromagnetic flow meter. Comparison between estimated blood flow by IPG and that obtained by volume displacement plethysmography made by Mohapatra and Arenson[26], Schraibman et a1[34] and Hill and Hope[11] has shown good correlation between the two estimations. Strain gauge plethysmography was observed to be giving consistently higher estimations than the impedance method. Though these studies validated the blood flow estimations by impedance plethysmography, they did not establish the genesis of the waveform.
Brook and Cooper[5] measured the impedance of the left index finger before and after the injection of normal saline into the brachial artery. Injection of saline, a better conductor than blood, produced sudden decrease in the impedance on reaching the finger. On the other hand, same volume of glucose, a poorer conductor than blood, when injected produced an increase in the impedance of the finger. This suggested that impedance plethysmograph responded to the changes in ionic concentration within the vascular compartment and the impedance change recorded is not due to the change in contact impedance at body electrode interface. Anderson et al[1] made injections of normal saline and non-conducting silicon oil into the calves of human cadavers. Injections (2ml) were made at two depths, just under the skin and near the centre of the muscle mass, at each of the 7 to 9 longitudinal positions. The results on 20 limbs from 12 cadavers confirmed that only injections of saline into region between sensing electrodes caused significant impedance change. They concluded that impedance plethysmography reflected changes in conductance confined to the area between the sensing electrodes. Brown et al[6] applied three sensing electrodes around the human calf at a separation of 100 min. Impedance changes were recorded between the three combinations of sensing electrodes. It was observed that the summation of the impedance changes between the upper and lower pair of electrodes was equal to that recorded between the uppermost and the lowermost electrodes. This experiment conclusively indicated that changes in contact impedance were much smaller than the changes in tissue impedance. Shimazu et al[35] in their experiment, have measured the admittance (reciprocal of impedance) of a conductivity cell. They have shown that the changes in the admittance produced by blood flow in the human finger dipped in the electrolyte of the conductivity cell vanishes when the conductivity of the electrolyte is equal to that of the blood. Yamakoshi et a1[42] have also used similar set up for non-invasive determination of haematocrit percentage. These studies provide experimental verification for the parallel conductor theory of Jan Nyboer, establish impedance blood-volume relationship and rule out genesis of IPG waveform from changes in contact impedance.
First time derivative of the impedance recorded in Kubicek's neck-abdomen configuration, commonly known as impedance cardiogram (ICG) is a precisely defined waveform in the time domain. Fig. 1 schematically represents the ICG waveform and its relationship to the cardiac cycle. The normal sinus rhythm ICG is comprised of three distinct waves, A-wave, C-wave and O-wave. The A-wave corresponds to the atrial systole and can occur independently of the C and O waves and is therefore absent before premature ventricular contraction[19],[38]. It appears 40-100 ms after the P-wave of ECG, usually as an increase in the impedance (negative wave in ICG)[23]. The largest deflection in a normal impedance cardiogram, the C-wave, leaves the base line at point B, reaches the peak at point C and ends, below the base line at point X. The B point has been observed to be synchronous with maximal deflection of the first heart sound at the apex of the heart[23]. Simultaneous recording of M-mode echo-cardiogram and impedance cardiogram has shown B point to be synchronous with the opening of the aortic valve[25],[32]. The C-point has been observed to occur very close in time to the peak of ascending aortic blood flow as measured with a cuffed electromagnetic flow probe placed around the root of the aorta[21]. The X point has been shown to be synchronous with aortic component of second heart sound[23] and also with the closure of aortic valve in M-mode echo-cardiogram[25],[32]. The C-wave is followed by a second major deflection called the O-wave. This wave begins at the X point, reaches the summit at O point and ends near the base line. Sometimes Y-wave is present in the upstroke of O-wave and is identified as a significant change in slope between X and O points. Similarly, another significant change in slope occurs in the down stroke of O-wave and is called as Z-wave. The Y point has been observed to be synchronous with pulmonary component of second heart sound. The O-point in time domain is coincident with the mitral opening snap of phonocardiogram[23] and also with mitral valve opening E-point) of M-mode echo-cardiogram[25],[33]. The Z-point is observed to be synchronous with the third heart sound representing the end of rapid filling phase[23]. Thus, impedance cardiogram is a precisely defined waveform in the time domain and represents most of the important events of the cardiac cycle. Sometimes the C-wave is preceded by a small deflection between A and B points. The peak of this deflection is named as I-point. However, no cardiac event has been associated with this point.
The investigation into the origin of impedance variations in the thorax is more complicated than that in the extremities due to simultaneous blood volume changes taking place in different compartments of the thoracic cavity. For example, simultaneous contraction of both the ventricles and the subsequent rapid ejection of blood produces many changes within the thorax. There is a decrease in the ventricular blood volume whereas there is an increase in the blood volume in the aortic arch, pulmonary arteries and the pulmonary capillaries. Thus, some cardiac events contribute an increase in the thoracic impedance while others contribute a decrease and therefore the impedance cardiogram recorded by surface electrodes is a resultant of number of events. Since the impedance measurement is derived from the surface potentials, any change in physical dimension and/or spatial rearrangement of the volume conductors of different resistivities causes a redistribution of current density, which appears as a change in the impedance between the measuring electrodes. It is, however, important to understand whether any specific cardiac event is predominantly reflected in the particular segment of the thoracic impedance cardiogram and under what circumstances this is superceded by changes from other sources. Bonjer et a1[4] by examining the results of electrically isolating the heart and lungs of a living dog and rhythmic perfusion of the systemic and the pulmonary circulation in the dead animal, have shown that the impedance changes principally occur from the expulsion of blood during ventricular contraction with negligible contribution from volume changes of the heart. By injecting fresh room temperature blood into isolated segments of the ascending and descending aorta, the right atrium and left ventricle in dogs, Kubicek et al[21] have shown a decrease in the thoracic electrical impedance and that for a fixed volume change the impedance change associated with it is greater in the aortic segments than in the ventricles. They have also shown a close correlation between the peak value of C-wave of ICG and the peak value of aortic blood flow in the ascending aorta measured with a cuffed electromagnetic flow meter. Contrarily, Geddes and Baker 8 have shown significant contribution from right ventricle by observing thoracic impedance changes following injection of saline into the right and left ventricles of dogs. By perfusing the aorta and the pulmonary artery of dogs with a control led sinusoidal or pulsatile flow of blood, Ito et al[15] have shown that the main component of ICG originates from the systemic rather than the pulmonary blood flow. In the modelling of Girling and Imrie[9], they have shown that the modelled ?Z(t) waveform is in agreement with the observed waveform, when contribution from aorta, chambers of the heart and vena-cava are included. Exclusion of any of these contributions has been shown to affect the modelled waveform significantly. However, exclusion of aortic contribution has been seen to produce marked variation in the modelled waveform. Similarly, Yamakoshi et al[41] have simulated on a digital computer the relationship between the actual blood flow into the aorta, the transthoracic impedance variation and the stroke volume from Kubicek's equation using a simple vascular tube model. The effect of varying Zo, the aortic elasticity, the peripheral resistance of the vascular bed and the velocity of the pressure pulse in the aorta was also investigated. They have shown that the maximum flow velocity, the left vertricular ejection time and the rise and fall times of the aortic inflow were the most important factors determining the magnitude of the calculated stroke volume. Good correlation between stroke volume estimated by Kubicek's method and that obtained by other standard techniques like dye dilution and Fick method[2],[7],[13],[20],[24],[27],[30],[36],[37], though an overestimate by ICG, supports genesis of C-wave of ICG from left ventricular ejection. In the experiments of Witose and Kottke[40] and Mohapatra[25], the transient occlusion of pulmonary flow has not affected the ICG significantly whereas that of the aorta reduced the amplitude of C-wave by more than 50%. A controversial increase of 68% in the amplitude of impedance cardiogram has been recorded, when the ventricles were isolated by an insulating sheet, in order to account for the remaining amplitude of the C-wave. The fact, that the peak of the putsatile outflow, measured at the aortic root by the electromagnetic flow meter [21] and the PH wave of ballistocardiogram[12] coincide with the peak of dZMt, confirms that peak occurs at the time of maximum ejection rate. The strong correlation between the peak dZ/dt height and the peak aortic flow, again emphasizes that the systolic portion of the Mdt signal, the C-wave, mainly reflects the rate of change of ventricular ejection. High correlation (0.90) between aortic regurgitation fraction obtained from ICG and that from angiography and also between area under systolic wave and angiographic stroke volume (0.96) further confirm the origin of C-wave from left ventricular ejection[32]. Very little is known about the 0wave and the A-wave, except that changes in these waves are observed in head-stand position. An increase in the amplitude of O-wave has also been observed in some of the patients with mitral stenosis[22], mitral regurgitation[18],[29],[33] and myocardial infarction[31] and on exposure of subjects to high altitudes[3]. Similarly an increase in the amplitude of A-wave is recorded, when subjects are suddenly exposed to high altitudes[3]. An induced augmentation of venous return during the head-stand is likely to result in an increased filling of the great veins and the atria. The inability of the ventricles, to handle this unusual amount of blood, probably causes an increase in the amplitude of O-wave. Good correlation between the mitral regurgitating farction obtained by angiography and that obtained from the area under the O-wave and C-wave of ICG, in patients with mitral regurgitation, further confirms the origin of the O-wave from atrio-ventricular transport. It is further supported by the increase in amplitude of O-wave in patients with tricuspid regurgitation. On the basis of studies conducted by us in more than 1000 patients with various cardaic disorders, we feel that maximum contribution to the ICG waveform comes from aorta and the venacava as they are parallel to the axis of impedance measurement. The contribution from various chambers of the heart is expected to be smaller as they do not cover the entire length of the measurement axis. Pulmonary arteries, being oblique to the axis of measurement, have small projections along the axis of measurement. Right ventricular output into the lungs also has negligible contribution as the blood is only displaced from one compartment to another in the thoracic cavity and mainly causes spatial rearrangement of the conductor. By making impedance measurement along the horizontal axis, we have shown that the waveform thus obtained has marked variations from alteration in pulmonary circulation[16],[17]. These observations confirm the genesis of conventional impedance cardiogram from the haemodynamic changes in the aorta and the venacava.
The authors are thankful to Dr. (Mrs) P Pai, Dean, Seth GS Medical College & King Edward Memorial Hospital, Shri MK Gupta, Assoc. Director, E & I Group, BARC, Shri BR Bairi, Head, Electronics Division, BARC and Shri KR Gopalakrishna, Head, Nuclear Instrumentation Section, BARC for encouraging this work.
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