reference is now made to the drawings wherein like numerals refer to like parts throughout . it is noted that while the invention is described herein in terms of an apparatus and method for determining cardiac output suitable for use on the thorax of a human subject , the invention may also conceivably be embodied or adapted to monitor cardiac output at other locations on the human body , as well as monitoring cardiac output on other warm - blooded species . all such adaptations and alternate embodiments are considered to fall within the scope of the claims appended hereto . referring now to fig3 a - 5 , the general methodology of measuring cardiac output in a living subject according to the invention is described . as previously discussed , the thoracic impedance z t ( t ) of a living subject may be modeled as comprising a constant impedance , z o , and time - varying impedance , δz ( t ). according to the well known “ parallel - column ” model of the thorax , this change in thoracic impedance , δz ( t ), is related to the pulsatile blood volume change . in this model , illustrated in the form of a schematic diagram in fig1 herein , effectively constant tissue impedances such as bone , muscle , and fat are modeled as a conducting volume z o 102 in parallel with the pulsatile impedance of the blood δz ( t ) 104 . this second impedance 104 is a time - varying fluid column with resistivity , ρ , cylindrical length , l , and a time - varying cross - sectional area that oscillates between zero and a value a , the latter which correlates to the stroke volume v . when the pulsatile volume is at a minimum in the cardiac cycle , all the conducting tissues and fluids are represented by z o . during the cardiac cycle , the cylinder cross - sectional area increases from zero until the cylinder &# 39 ; s volume equals the blood volume change . because z o is much greater than δz ( t ), the relationship of eqn . ( 2 ) holds : sv = ρ  ( l 2 z 0 2 )  vet   z  ( t )  t min , ( 2 ) where l is the distance between the measurement electrodes in cm ( fig3 a ), vet is the ventricular ejection time in seconds , and  z  ( t )  t min is the magnitude of the largest negative derivative of the impedance change occurring during systole in ohms / s . often , the impedance derivative 400 is purposely inverted as shown in fig4 so that the original negative minimum change will appear as a positive maximum 402 ,  z  ( t )  t max , the ventricular ejection time ( vet ) is estimated from features in the impedance waveform , which is obtained from the measurement terminals of the electrode arrays 302 , 304 , 306 , 308 placed on various locations of the subject &# 39 ; s thorax as illustrated in fig3 a and 3 b . in the present embodiment , a value of 150 ohm - cm is used for the resistivity of the blood , although it will be recognized that other values may be substituted as appropriate . it is noted that the description of the volume of participating tissue may be modified . rather than model the thorax as a cylinder as shown in fig1 above , the thorax may instead be modeled as a truncated cone ( as first described by sramek and bernstein ). this approach results in a modified stroke volume calculation as in eqn . ( 3 ): sv = l 3 4 . 25  z 0  vet   z  ( t )  t min . ( 3 ) with either of the two aforementioned approaches ( i . e ., cylindrical or truncated cone ), the pulsatile impedance is estimated using ohm &# 39 ; s law , which is well known in the electrical arts . specifically , current from a constant current source , i t ( t ), is applied , and the resulting voltage , v t ( t ), is measured in order to calculate the ratio of eqn . ( 4 ): z t  ( t ) = v t  ( t ) i t  ( t ) . ( 4 ) in the selected frequency range ( i . e ., 68 khz ), the typical impedance associated with a human subject &# 39 ; s skin is 2 to 10 times the value of the underlying thoracic impedance z t ( t ). to aid in eliminating the contribution from skin and tissue impedance , the present invention uses at least two , and typically four electrode arrays 302 , 304 , 306 , 308 for measurement , as shown in fig3 a . the physical construction and these electrode arrays is described in detail with reference to fig7 a - 8 herein . in a simple application , one electrode array 302 comprising a stimulation electrode terminal 310 and a measurement electrode terminal 312 is applied above the thorax 300 of the subject , while a second electrode array 304 ( having stimulation electrode terminal 314 and measurement electrode terminal 316 ) is applied below the thorax 300 . the ac current from the current source is supplied to the stimulation electrode terminals 310 , 314 . as shown in fig3 b , current flows from each stimulation electrode terminal 310 , 314 through each constant skin impedance , z sk1 , or z sk4 , each constant body tissue impedance , z b1 or z b1 , and each constant skin impedance , z sk2 or z sk3 , to each measurement electrode terminal 312 , 316 . the voltages at the measurement electrode terminals 312 , 316 are measured and input to a differential amplifier to obtain the differential voltage , v t ( t ). the desired thoracic impedance , z t ( t ), is then obtained using the relationship of eqn . ( 4 ). as shown in fig3 a , two sets of electrode arrays may advantageously be used to monitor the impedance associated with the left and right portion of the thorax 300 in the present invention . when eight electrode terminals ( four arrays 302 , 304 , 306 , 308 ) are used in this manner , the four measurement arrays are also used to obtain an electrocardiogram ( ecg ), based on one of four vectors modified from lead i , i , iii , or iv . the resulting electrocardiograms are based on the original lead configurations , but are not of diagnostic quality . regardless of the modified lead configuration used , the q wave of the ecg qrs interval is used to determine the heart rate and to trigger measurements of vet within the  z  ( t )  t fig5 illustrates the logical flow of the method of measuring cardiac output according to the invention . as shown in fig5 the method 500 generally comprises first providing a plurality of electrode “ arrays ” of the type previously described herein per step 502 . the electrode arrays are positioned at predetermined locations above and below the thoracic cavity per step 504 , as illustrated in fig3 a herein . in one embodiment of the method , these locations are chosen to be on the right and left sides of the abdomen of the subject , and the right and left sides of the neck . these locations , with prior art band electrodes , were first used by kubicek . other locations and / or combinations of arrays may be substituted with equal success . next , a substantially constant ac current is generated in step 506 , and the current applied to the stimulation electrode terminal 310 , 314 of each of the electrode arrays in step 508 . the voltage generated at the measurement electrode terminal 312 , 316 of each electrode array is next measured in step 510 . as previously discussed , this voltage is generally reduced from that applied to the stimulation electrode by virtue of the impedance of , inter alia , the thoracic cavity . note that the measured voltage may be absolute , or relative ( i . e ., a differential voltage ) as desired . next , in step 512 , the cardiac stroke volume from the measured voltage , using for example the relationship of eqn . ( 3 ) above . cardiac rate ( step 514 ) is also determined by using the measurement electrodes to sense the ecg potentials generated by the heart of the subject . lastly , in step 516 , cardiac output is determined based on the stroke volume determined in step 512 and the cardiac rate in step 514 using the relationship of eqn . 1 above . referring now to fig6 the apparatus for measuring cardiac output using the above - described technique is disclosed . in addition to the four electrode arrays 302 , 304 , 306 , 308 previously discussed , the system 600 generally comprises an alternating current ( ac ) current source 604 capable of generating a substantially constant current , a plurality of electrical leads in the form of a multi - ended lead assembly 606 for connecting the instrument monitor 607 to the individual terminals of the electrode arrays 302 , 304 , 306 , 308 , a processor 608 with associated algorithms capable of running thereon for performing analysis of the signals measured from the measurement terminals , data and program memory 609 , 610 in data communication with the processor 608 for storing and retrieving program instructions and data ; an i / o interface 611 ( including analog - to - digital converter ) for interfacing data between the measurement electrodes and the processor 608 ; a mass storage device 612 in data communication with the processor for storing and retrieving data ; a display device 614 ( with associated display driver , not shown ) for providing an output display to the system operator , and an input device 616 for receiving input from the operator . it will be recognized that the processor 608 , memory 609 , 610 , i / o interface 611 , mass storage device 612 , display device 614 , and input device 616 ( collectively comprising the instrument monitor 607 ) may be embodied in any variety of forms , such as a personal computer ( pc ), hand - held computer , or other computing device . the construction and operation of such devices is well known in the art , and accordingly is not described further herein . the applied current derived from the current source 604 is a 70 khz sine wave of approximately 2 . 5 ma peak - to - peak . the measured voltage associated with the aforementioned sine wave is on the order of 75 mv peak - to - peak . these values are chosen to advantageously minimize electric shock hazard , although it will be appreciated that other frequencies , currents , or voltages may be substituted . the construction and operation of ac current sources is well known in the electronic arts , and accordingly is not described further herein . the electrode lead assembly 606 of the illustrated embodiment contains a ten wire assembly ( two wires are left unused ) that branches to eight individual connectors 606 a - h . the conductors 610 a - h of the lead assembly are fashioned from electrically conductive material such as copper or aluminum , and are insulated using a polymer - based insulation having the desired dielectric strength as is well known in the electrical arts . the length of the conductors may also be controlled so as to match the impedance of each individual conductor to that of the others within the assembly 606 . using one of four modified lead configurations , the body surface potential is measured between two measurement electrodes . this time - varying voltage reflects the electrical activity of the heart , and contains one qrs interval per cardiac cycle . the biopotential is analyzed to identify each qrs complex . the frequency of qrs complexes determines the heart rate . the q wave within the qrs complex is then used to trigger identification of vet within the  z  ( t )  t waveform , as the opening of the aortic valve ( the beginning of vet ) occurs after the appearance of the q wave . referring now to fig7 a - 7 c , the electrode arrays 302 , 304 , 306 , 308 of the invention are described in detail . as illustrated in fig7 a , each array comprises a flexible substrate 704 having a plurality of apertures 706 , 708 formed therein . in the illustrated embodiment , two terminals 310 , 312 are disposed through the apertures such that the top portions 716 , 718 of the terminals project above the plane of the substrate 704 . the two terminals 310 , 312 comprise a stimulation terminal 310 and measurement terminal 312 as previously described with respect to fig3 a . the stimulation terminal 310 is used to apply the potential necessary to generate the current flowing through the thoracic cavity of the subject . it will be noted that despite designation of one terminal as a “ stimulation terminal ” and one as a “ measurement ” terminal , the role of these terminals may be reversed if desired , since they are functionally and physically identical but for the potential applied thereto ( or measured therefrom ). it is noted that the asymmetric shape of the substrate 704 of the embodiment of fig7 a - 7 c may be used to assist the clinician in rapidly determing which electrode is the stimulation electrode and which the measurement electrode , such as by assigning a convention that the end of the array having a given shape always contains the stimulation electrode . additionally , the substrate may be shaped to adapt to certain physical features of the patient , such as by using a substrate having a broader width so as to better conform to the generally cylindrical shape of the subject &# 39 ; s neck . any number of different substrate shapes may be employed ; fig7 d illustrates one such alternative shape . as shown in fig7 a - 7 c , the terminals 310 , 312 are firmly held in place within the substrate 704 at a predetermined distance 705 by a mounting element 707 or any one of a variety of other constructions as will be described in greater detail below . the distance ( measured centerline - to - centerline on the terminals 310 , 312 ) is approximately 5 cm in the embodiment of fig7 a , although it will be recognized that other distances may be substituted . desired distances may be determined through experimentation , anecdotal observations , calculations , or any other suitable method ; however , experimental evidence obtained by the applicant herein indicates that a distance of 5 cm is optimal for impedance cardiography measurements . the substrate 704 in the embodiment of fig7 a is formed from a polyethylene foam , although other materials such as cloth or vinyl may be substituted . the polyethylene foam is chosen for its compliance and flexibility , thereby allowing it to conform somewhat to the contours of the subject &# 39 ; s anatomy , while still maintaining sufficient rigidity for maintaining the terminals 312 , 314 in the desired position and orientation . as shown in fig7 b , the terminals 310 , 312 of each array comprise a generally cylindrical shaped sidewall portion 720 having a first diameter 722 , and a top portion 724 having a second diameter 726 , the second diameter 726 being greater than the first diameter 722 in order to assist in retaining a connector mated to the terminal 310 , 312 as described in greater detail below . the outer wall 721 of the sidewall portion 720 is essentially vertical in orientation ( i . e ., parallel to the central axis 725 of the terminal 310 , 312 ), while the top portion is progressively rounded as shown . the terminals may be manufactured from an extruded metal such nickel , with a coating of brass , or may be molded from carbon . alternatively , the terminals may be molded of plastic , and coated with a metal such as gold or impregnated with carbon . the extruded metal possesses the advantage of low cost , while the molded plastic impregnated with carbon possesses the advantage of radiolucency . a terminal molded of plastic and coated with gold may possess low noise artifact . the terminals 310 , 312 of the electrode array comprise a two piece construction , having an upper terminal element 730 and a lower terminal element 732 as shown in fig7 a and 7 b . the post 734 of the lower terminal element 732 is adapted to be frictionally received within the cavity 736 of the upper terminal element when the two components are mated . in this fashion , the upper and lower elements 730 , 732 form a single unit when assembled , with the mounting element 707 being frictionally held or “ pinched ” between the lower surface 740 of the upper element 730 and the upper surface 742 of the lower element 732 . the post 734 of the lower element perforates the mounting element 707 , or alternatively penetrates through a pre - existing aperture 738 formed therein . the lower elements 730 , 732 of the electrode array terminals 310 , 312 are coated with ag / agcl , although other materials with the desirable mechanical and electrochemical properties such as zinc chloride may be used if desired . the electrolytic element 750 of each electrode array comprises an electrolytic gel of the type well known in the bio - electrical arts ; in the present embodiment , the gel comprises an ultraviolet ( uv ) cured potassium chloride ( kcl ) gel , although it will be recognized that other types of compounds or materials may be used . uv curing of the gel allows the element 750 to have a more solidified consistency and improved mechanical properties , thereby preventing excessive spreading or thinning of the element when the array is applied to the subject while still maintaining its overall adhesiveness and electrolytic properties . as shown in fig7 b and 7 c , the element 750 is sized so as to encompass the edges 752 of the respective aperture 706 , 708 in the substrate 704 over which it is placed when assembled , although other configurations may be used . the top portion 755 of the element 750 fits at least partially within the aperture 706 , 708 and conforms substantially thereto , thereby effecting contact with the bottom surface 760 of the bottom terminal element 732 . in this way , ions are passed between the skin of the subject and the terminals of the array via the gel element 750 . the gel also provides for adhesion of the array to the skin of the subject , although the array of the present embodiment also includes a separate adhesive 762 which is applied to the bottom surface of the substrate 704 , as shown in fig7 c . since the placement of the electrolytic element 750 with respect to the terminals 310 , 312 of the array may in certain cases affect the ultimate measurements of cardiac output obtained with the system , the gel of the element 750 is advantageously placed in the embodiment of fig7 a - c so as to be symmetric with respect to the terminal 310 , 312 . it will be recognized , however , that the element ( s ) 750 may also be placed so as to produce certain desired electrolytic conditions . similarly , the element 750 may be split into two or more component parts if desired . furthermore , it is noted that while the embodiment of fig7 a - c employs two fixed terminals that are effectively immovable within the substrate , means for allowing adjustment or change of the relative position of the terminals may be substituted . for example , as illustrated in fig8 a terminal array having three terminal posts may be used , the second post 802 being spaced a first distance 804 from the first post 806 , and the third post 810 being spaced a second distance 808 from the first post 806 , such that the clinician can select one of two terminal spacings as desired . as illustrated in fig9 each electrode lead assembly connector 606 a - h is designed to mitigate the downward force required to mate the connector with its respective electrode array terminal . specifically , each connector 606 a - h contains two spring - biased conductive jaws 902 that are spread apart by the cam surface 904 of an actuator button 906 disposed on the front 907 of the connector body 908 . the connector jaws 902 and bias mechanism are designed to allow the upper and sidewall portions 724 , 720 of the electrode terminal 310 , 312 ( fig7 b ) to be received within the recess 910 of the jaws 902 when the button 906 is fully depressed . in this fashion , effectively no downward force is required to engage the connector to its respective terminal . the jaws 902 are contoured to engage substantially the entire surface of the sidewall portion 720 of the terminal when the actuator button 906 is released . since the sidewall portion 720 of the terminal is effectively circular in cross - section , the connector may advantageously rotate around the axis of the terminal 310 , 312 when lateral tension is applied to the conductor attached to that connector . u . s . pat . no . 5 , 895 , 298 issued apr . 20 , 1999 , entitled “ dc biopotential electrode connector and connector condition sensor ,” and incorporated herein by reference in its entirety , describes a bias jaw electrical connector of the type referenced above in greater detail . when used with the four two - terminal electrode arrays 302 , 304 , 306 , 308 shown in fig3 a , each connector 606 a - h is fastened to one of the two terminals 310 , 312 of an electrode array . the 68 khz constant current is applied from the current source to four electrode terminals ( i . e ., one terminal per array ). hence , complete circuits are formed between the current source and the i / o device 611 of the system 600 via the electrical conductors and connectors associated with the stimulation electrode terminals , the stimulation electrode terminals themselves , the thorax of the subject , the measurement terminals , and the electrical conductors and connectors associated with the measurement terminals . referring now to fig1 , the method of evaluating the electrical continuity of one or more leads within the system is described . note that while the following description is based on the two - terminal array configuration ( fig7 a - 7 c ) and the use of four arrays as shown in fig3 a , the method may be applied to many alternate configurations with equal success . first , in step 1002 , the electrode arrays are disposed on the skin of the subject . the position at which the electrode arrays are disposed on the subject are measured in relation to the thoracic cavity as illustrated in fig3 a , or alternatively may be inferred by the weight and height of the subject . next , a current is generated between the stimulation electrodes and the measurement electrodes of the respective arrays in step 1004 . as previously discussed , the current passes through at least a portion of the subject &# 39 ; s thoracic cavity , encountering a time - variant impedance therein . an impedance waveform is then measured from two or more of the measurement terminals of the arrays in step 1006 . the waveforms comprise measurements of impedance as a function of time , which is well known in the cardiographic arts . these measured waveforms are then compared to one another in step 1008 to detect changes or variations between them . in the present embodiment , two waveforms are differenced by way of a simple differencing algorithm resident on the processor 608 of the system 600 ( fig6 ), although it will be recognized that other approaches may be used . for example , the base impedance may be calculated for the left and right sides . the larger base impedance may then be subtracted from the smaller base impedance , with this difference then divided by the smaller impedance . the resulting percentage ratio , when greater than a predetermined threshold value , may represent the presence of detached or loose electrodes . while some variation between the waveforms is normal , significant variations are indicative of either a degraded electrical connection , such as between the electrode array terminal and its respective connector , or between the electrolytic gel and the skin of the patient , or even the gel and the terminal of the array or between the cable and connector . a threshold value is determined and set by the operator of the system in step 1010 such that when the threshold “ difference ” is exceeded as determined by the aforementioned algorithm ( step 1012 ), the operator will be alerted to the degraded condition such as by a visual or audible alarm in step 1014 . it is noted that the use of the multi - terminal electrode arrays having predetermined and substantially equal terminal spacing as previously described allows such comparisons between electrode waveforms to be made ; errors resulting from uncontrolled spacing of the terminals are effectively eliminated . using prior art electrodes , the aforementioned method would be largely ineffective , since these error sources would force the threshold value to be set artificially high , thereby potentially masking conditions of degraded electrical continuity which could affect the ultimate accuracy of and cardiac output estimation made by the system . it will be recognized that while certain aspects of the invention have been described in terms of a specific sequence of steps of a method , these descriptions are only illustrative of the broader methods of the invention , and may be modified as required by the particular application . certain steps may be rendered unnecessary or optional under certain circumstances . additionally , certain steps or functionality may be added to the disclosed embodiments , or the order of performance of two or more steps permuted . all such variations are considered to be encompassed within the invention disclosed and claimed herein . while the above detailed description has shown , described , and pointed out novel features of the invention as applied to various embodiments , it will be understood that various omissions , substitutions , and changes in the form and details of the device or process illustrated may be made by those skilled in the art without departing from the invention . the foregoing description is of the best mode presently contemplated of carrying out the invention . this description is in no way meant to be limiting , but rather should be taken as illustrative of the general principles of the invention . the scope of the invention should be determined with reference to the claims .
an improved apparatus and method for determining the cardiac output of a living subject . the improved apparatus generally comprises one or more electrode assemblies or patches affixed to the skin of the subject in the vicinity of the thoracic cavity . the terminals of each electrode patch are in contact with an electrolytic gel , and are spaced a predetermined distance from one another within the patch . this predetermined spacing allows for more consistent measurements , and also allows for the detection of a loss of electrical continuity between the terminals of the patch and their associated electrical connectors in the clinical environment . the method generally comprises generating and passing a stimulation current through the terminals and the thoracic cavity of the subject , and measuring the impedance as a function of time . this impedance is used to determine cardiac muscle stroke volume , which is then used in conjunction with the subject &# 39 ; s cardiac rate to determine cardiac output . a method of detecting a loss of electrical continuity in one or more of the terminals of the electrode patch is also disclosed .