J. Biomedical Science and Engineering, 2009, 2, 491-498
doi: 10.4236/jbise.2009.27071 Published Online November 2009 (http://www.SciRP.org/journal/jbise/
Published Online November 2009 in SciRes. http://www.scirp.org/journal/jbise
Electrocardiographic interference and conductance volume
Simon P. McGuirk1,2, Dan Ewert3, David J. Barron1, John H. Coote2
1Department of Cardiac Surgery, Birmingham Children’s Hospital, Birmingham, UK; 2School of Clinical and Experimental Medicine,
University of Birmingham, Birmingham, UK; 3Department of Electrical and Computer Engineering, North Dakota State University,
Fargo, USA.
Email: simon.mcguirk@nhs.net
Received 1 February 2009; revised 1 July 2009; accepted 16 July 2009.
The conductance catheter technique enables con-
tinuous ventricular volume measurements based on
the electrical conductance of blood within the ven-
tricular cavity. However, ventricular excitation also
produces a measurable electrical signal within the
ventricular cavity. This study was undertaken to in-
vestigate the relationship between the ventricular
electrogram and conductance volume measurements
in a physical model of the left ventricle without par-
allel conductance. The ventricular electrogram was
simulated with an ECG signal, ECGinput connected to
two ring electrodes within the model ventricle. Con-
ductance volume measurements were made with and
without ECGinput. The difference between these mea-
surements, GECG(t), represented the conductance
volume due to ECGinput. GECG(t) varied as a function
of the first-derivative of ECGinput with respect to time
(r2=0.92, P<0.001). GECG(t), This primarily affected
volume measurements during ventricular depolarisa-
tion; during this phase the volume measurement er-
ror varied widely between –12% and +9%. As a re-
sult, end-diastole could not be reliably identified on
the pressure-volume loop. The accuracy of conduc-
tance volume measurements during late diastole and
early isovolumic contraction are substantially af-
fected by the ventricular electrogram. This may re-
sult in a significant error in end-diastolic volume es-
timates, which has important implications for the
quantitative assessment of ventricular function in-
cluding, in particular, the assessment of chamber
Keywords: Volume Measurement; Conductance Catheter;
Electrocardiogram; Ventricular Electrogram
The assessment of ventricular function is fundamentally
important for the evaluation of patients with known or
suspected heart disease. Analysis of left ventricular (LV)
volume in the time and pressure domains allows systolic
and diastolic function to be separately quantified. The
conductance catheter technique was developed to con-
tinuously measure ventricular volume in real-time [1,2].
These measurements are recorded simultaneously with
intraventricular pressure measurements to provide in-
stantaneous pressure-volume data [3].
The conductance catheter technique is associated with
two, well known sources of error. Firstly, the current
density generated by the conductance catheter is not
uniformly distributed throughout the ventricular cavity
[4,5,6]. This results in a non-linear conductance-absolute
volume relationship [2,4,7]. Conductance volume meas-
urements must be corrected with a calibration coefficient,
α [8]. Secondly, the tissues and fluid surrounding the
ventricular cavity also contribute to the conductance
signal [2,8]. This results in an offset in the conductance-
absolute volume relationship, called parallel conduc-
tance. In practice, conductance volume measurements
are usually calibrated for parallel conductance using the
hypertonic saline method in order to derive accurate
ventricular volume measurements [8].
In our paediatric clinical experience, we have ob-
served that the pattern of LV volume measurements is
frequently abnormal. The conductance volume meas-
urements are characterised by a narrow upward spike
followed by a narrow downward spike during late dias-
tole without any commensurate change in LV pressure
(Figures 1A and 1B). This alters the shape of the pres-
sure-volume loop, with the loss of the normal lower
right-hand corner (Figure 1D). To our knowledge, this
abnormal conductance volume pattern has not previously
been described. However, we understand that similar
S. P. McGuirk. Lu et al. / J. Biomedical Science and Engineering 2 (2009) 491-498
Figure 1. Time-varying pressure (A), conductance volume
(B) and surface electrogram (C) signals obtained in a 6
year-old child with tricuspid atresia. The corresponding
pressure- conductance volume loop for the same child (D)
was developed by plotting the instantaneous pressure
against the corresponding conductance volume.
findings have been observed in other patient groups,
particularly in patients with a permanent pacemaker
[personal communication, Professor M.P. Frenneaux,
Department of Cardiovascular Sciences, University of
Birmingham, UK; Dr P. Steendijk, Department of Cardi-
ology, Leiden University Medical Center, The Nether-
We observed that the abnormal LV conductance vol-
ume measurements occurred synchronously with the
QRS complex on the surface electrocardiogram (ECG;
Figure 1C). We hypothesised that this abnormal LV
conductance measurement may represent the effect of
ventricular excitation on conductance volume measure-
ments, which is superimposed on the normal ventricular
volume cycle. This study was undertaken to examine the
relationship between the ventricular electrogram and
conductance volume measurements in a physical model
of the left ventricle without parallel conductance. In ad-
dition, we sought to determine how this relationship was
influenced by changes in electrical resistance across the
model ventricle.
2.1. Model Ventricle
This study used the physical model of the left ventricle
previously developed by this aboratory [9] and described
in the accompanying paper [10]. In summary, this con-
sisted of an ellipsoid latex balloon enclosed in a pressur-
ised Perspex chamber. The chamber was filled with dis-
tilled water and hydraulically pressurised with an in-
tra-aortic balloon pump (IABP; Datascope Medical Co.
Ltd, Huntingdon, UK) connected to two 25ml intra-aortic
balloon catheters in parallel. A patient simulator (Bioma
Research Inc, Quebec, Canada) was used to trigger the
IABP console at a predetermined rate (60 beats·min-1).
Inflation of the intra-aortic balloon catheters displaced
the stroke volume (SV, 50ml) from the model ventricle
through a 2/2-way solenoid valve (Bürkert GmbH,
Ingelfingen, Germany) into a calibrated measuring cyl-
inder at the top of the model ventricle. Deflation of the
IABP balloon catheters caused ventricular pressure to
fall, which allowed the latex balloon to refill. Electronic
circuitry was used to control the opening and closure
times of the solenoid valve in order to simulate different
contraction patterns.
The latex balloon was 13 cm in length and had a
maximal volume at atmospheric pressure of 500ml. The
balloon was filled with 385–500ml of buffered saline
solution (V) at room temperature. The saline concentra-
tion was varied between 0.18 – 1.57%. The resistivity (ρ;
conductivity-1) of these solutions was measured before
each test using a dedicated measuring cuvette (CD Ley-
com, Zoetermeer, The Netherlands). The resistivity
ranged between 37 ± 0.8 to 330 ± 0.5 cm.
Copyright © 2009 JBiSE
S. P. McGuirk et al. / J. Biomedical Science and Engineering 2 (2009) 491-498 493
Ventricular depolarisation was simulated using a fixed
output ECG signal, ECGinput, from the patient simulator
(maximum 150.4 ± 0.1 mV, minimum –35.8 ± 0.2 mV).
This was connected, via a resistor (200 ), to a dipole
within the latex balloon. This dipole consisted of two
copper ring electrodes (diameter 30mm, depth 5mm and
thickness 1mm) that were positioned perpendicular to
the long-axis of the balloon. The in-series resistor was
adjusted so that the signal range of the intracavitary
electrogram (ECGc) in the model ventricle was equiva-
lent to that observed in vivo (~1mV).
In vivo, the intracavitary electrogram primarily re-
flects the pattern of ventricular depolarisation and repo-
larisation in the endocardium of the ventricle [11]. The
position of the endocardium relative to the long-axis of
the ventricle will vary as the ventricular volume changes
during the cardiac cycle. This effect was simulated by
altering the distance between the two electrodes of the
dipole. The distance between the two electrodes (D) was
varied between 3 cm and 11.5 cm such that both elec-
trodes remained equidistant from the centre of the bal-
The principle of the conductance catheter technique for
measuring LV volume has been described elsewhere [8].
The details of the conductance catheter used in this study
are described in the accompanying paper [10]. The
catheter measured seven time-varying segmental con-
ductance signals, Gi(t). As parallel conductance is negli-
gible in our model, the total conductance volume, Q(t),
was determined using the following formula:
()[ ()]
QtLG t
where ρ is the blood resistivity and L is the in-
ter-electrode distance. The dimensionless calibration
coefficient, αV(t) was calculated from conductance vol-
ume measurements without the ECG signal by dividing
the conductance-derived volume measurement by the
absolute volume at either end-diastole or end-systole:
() or
We have previously demonstrated that the calibration
coefficient, αV(t) varies as a non-linear function of the
absolute ventricular volume [10]. We used this
non-linear αV(t)-volume relation to calibrate conductance
volume measurements:
() ()
Vt Qt
 (3)
where αVV is the αV(t)-volume relation [10].
Instantaneous pressure within the model ventricle was
measured using a high-fidelity solid-state micromano-
meter laterally positioned between electrodes 5 and 6
within the conductance catheter. This pressure signal
was amplified using a combined amplifier-interface unit
(PCU-2000; Millar Instruments, Houston, TX, USA) and
statically calibrated using a separate fluid-filled cathe-
ter-manometer system.
The cavitary electrogram, ECGc was measured as part
of the conductance catheter technique. The conductance
signal between electrodes 5 and 6 was measured, ampli-
fied and filtered using a second-order filter with a high
cut-off frequency (–3dB; 125 Hz) in order to derive the
ECGc [personal communication; CD Leycom, Zoeter-
meer, The Netherlands].
The conductance signal was measured either with,
, or without the ECG signal, . ECG inter-
ference, GECG(t) was calculated as the difference be-
tween these two conductance signals:
()() ()
ECG ii
GtGt Gt
The difference between calibrated volume measure-
ments made with, and without the ECG signal,
was expressed as a percentage of the
[() ()]
Vt Vt
 (5)
Analogue signals representing 7 segmental conductance sig-
nals, the pressure within the model ventricle, and both
ECGinput and ECGc were all digitised at 12-bit accuracy
and a sample frequency of 250 Hz. End-diastole and end-
systole were retrospectively identified. End diastole was de-
fined as the R wave on the ECG and end-systole was defined
as the point immediately prior to IABP circuit deflation.
The effect of intracavitary volume (V), resistivity of
the saline solution (ρ) and inter-electrode distance (D) on
the intracavitary electrogram and conductance volume
measurements were examined in turn. This involved a
series of experiments, in which one variable was altered
incrementally while the other two variables remained
unchanged. This process was repeated until the data
from the entire range was obtained (see above).
Each experiment was conducted under steady-state con-
ditions and data from 5 consecutive cycles were analysed.
The average within-experiment standard deviation was 0.71
ml and, at its worst, this represented <0.5% of the total
conductance volume. All subsequent analyses were there-
fore based on the average data from each experiment.
Copyright © 2009 JBiSE
S. P. McGuirk. Lu et al. / J. Biomedical Science and Engineering 2 (2009) 491-498
Data were analysed using SPSS for Windows (v12,
SPSS Inc., Chicago, Il, USA). Data are expressed as
mean ± SD and comparative analyses have been made
using the t-test. The relationship between the ECGinput,
ECGc and the ECG interference pattern was evaluated
by least squares linear regression based on fractional
polynomials of the data. The intracavitary volume,
resistivity of the saline solution and the inter-electrode
distance were included as covariables in the regres-
sion analyses. The coefficients of the linear regression
analyses, in both the overall and covariance analyses,
are expressed as mean ± standard error and a prob-
ability, P<0.05, was taken to represent statistical sig-
nificance. The “goodness of fit” of the prediction
equation was assessed as the square of the correlation
between dependent and significant independent vari-
7.1. Comparison between ECG Signal and
Cavitary Electrogram
The ECG signal, ECGinput, consisted of P, Q, R, S and T
deflections that resembled the normal lead II electrocar-
diogram (Figure 2A). The P wave was 84 ms in duration
with a peak of 6.2 mV (40 ms). The P wave represented
approximately half the PR interval (156 ms). The posi-
tive QRS complex had an overall duration of 84 ms,
with a maximum at 192 ms (150.4 mV) and two minima
at 164 ms (–26.4 mV) and 224 ms (–34.1 mV). The ST
segment was isoelectric (–8.2 mV) and 36 ms long. The
duration of the T wave was 236 ms, with a peak of 41.4
mV (408 ms). Finally, the QT interval and TP segment
were both isoelectric and 120 ms and 468 ms long, re-
The ECGc signal resembled the ECGinput signal turned
upside down, with an inverted P wave, an rSR wave and
an inverted T wave (Figure 2B). The overall relationship
between the two signals was best approximated by a
mathematical model in which the ECGc signal was in-
versely proportional to the ECGinput signal (r2=0.74,
() ()
 (6)
Although the ECGinput and ECGc signals were similar,
they were not identical There were differences in timing
and amplitude of the two signals. The ECGc P and S
wave minima and the R wave peak occurred either syn-
chronously or within one data point (i.e. 4 ms) of the
corresponding points on the ECGinput signal. By contrast,
the ECGc R wave peak and the T wave misnimum were
12 ms and 32 ms earlier than the corresponding points
on the ECGinput signal. The ECGc R wave peak was also
disproportionately pronounced compared to the cor-
responding S wave of the ECGinput signal. The ECGc
R-S wave ratio was –0.51 ± 0.03 whereas the ECGinput
S-R wave ratio was –0.17 ± 0.02 (P<0.001). In addi-
tion, the ECGc did not accurately reproduce the
isoelectric phases of the ECGinput signal. During the
PR, RT and TP segments, the ECGc signal was ini-
tially elevated and decreased progressively towards
the baseline signal.
In the covariance analyses, the intercept value (β0)
varied as a linear function of inter- electrode distance
(P<0.05), but was not affected by variation in the other
two factors. By contrast, the linear regression coefficient
(β1) varied as the inverse function of intracavitary vol-
ume and as a direct function of inter-electrode distance
and resistivity of the solution (all P<0.05).
When these effects are combined, the relationship
between ECGc and ECGinput was influenced by the in-
ter-electrode distance (intercept value) and by the total
resistance of the volume conductor (Eq. 7; r2=0.65):
23 4
()( )()
 
 
where β2 = 3.66 ± 0.01; β3 = 7.29·10-3 ± 1.53·10-3 and β4
= –7.00·10-3 ± 0.06·10-3 (P<0.05 for each coefficient).
7.2. Comparison between ECG Interference and
ECG Input Signals
The ECG interference signal, GECG(t) was characterised
by a low amplitude biphasic P wave; a high amplitude
equiphasic qRSr complex; and a low-amplitude biphasic
T wave (Figure 2C). Each phase of the GECG(t) signal
was synchronous with the P wave, QRS complex and T
wave of the ECG signal, respectively.
The amplitude of the GECG(t) has been described as a
percentage of the maximum GECG(t) signal from the
isoelectric line. The GECG(t) P wave had a sine wave-
like appearance with an initial upward deflection im-
mediately followed by a downward deflection of com-
parable duration and amplitude. The maximum and
minimum GECG(t) P wave signals were 4.0 ± 2.6% (28
ms) above and 4.4 ± 2.6% (64 ms) below the isoelectric
line. The spiked wave GECG(t) qRSr complex had two
maxima at 180 ms (R wave; 100 ± 3%) and 232 ms (r
wave; 22 ± 3%) and two minima at 160 ms (q wave;
–24 ± 3%) and 204 ms (S wave; –107 ± 3%). The
GECG(t) q and R waves occurred 32 ms and 12 ms be-
fore the ECGinput R wave whereas the GECG(t) S and r
waves occurred 12 ms and 40 ms after the ECGinput R
wave. The GECG(t) T wave had a similar overall appear-
ance to the GECG(t) P wave with an initial upward deflec-
tion immediately followed by an equivalent downward
deflection. The maximum and minimum GECG(t) T wave
Copyright © 2009 JBiSE
S. P. McGuirk et al. / J. Biomedical Science and Engineering 2 (2009) 491-498 495
Figure 2. The ECG signal (ECGinput; A), cavitary electrogram
(ECGc; B) and ECG interference (GECG(t); C) signals versus
signals were 4.4 ± 2.6% (352 ms) and 4.8 ± 2.6% (452 ms)
below the isoelectric line and occurred 80 ms and 180 ms
after the start of the ECGinput T wave, respectively.
The relationship between GECG(t) and ECGinput was well
approximated (r2=0.92, P<0.001) by a regression equa-
tion in which the interference signal varied proportionally
to the first-derivative of ECGinput with respect to time:
() ()
ECG input
Gt dECGt
 (8)
In the covariance analyses, the intercept (β0) values
varied as a linear function of the interelectrode distance
(P<0.05), but was not affected by variation in the other
two factors. By contrast, the linear regression coeffi-
cients (β1), varied as a function of the intracavitary vol-
ume; and as the inverse function of both the in-
ter-electrode distance and the resistivity of the solution
(all P<0.05).
Overall, the relationship between GECG(t) and the first
derivative of ECGinput varied as a function of the in-
ter-electrode distance and the conductivity of the volume
conductor (Eq. 9; r2=0.88):
() ()()
ECG input
where β0 = 1.49·10-3 ± 0.39·10-3 and β1 = 6.07·10-3 ±
0.38·10-3 (P<0.05 for both coefficients).
7.3. ECG Interference and Calibrated
Conductance Volume Measurements
For the purposes of this simulation, the inter-electrode
distance was assumed to change in accordance with the
instantaneous volume within the latex balloon. The in-
ter-electrode distance was estimated as the maximal
short-axis diameter of the spheroid, which varied
from7.0 cm (VES = 335 ml) to 8.6 cm (VED = 500 ml).
Calibrated conductance volume measurements with the
ECG signal, were compared against synchro-
nous calibrated conductance volume measurements
without the ECG signal, .
A representative example of calibrated conductance
volume measurements with and without the ECG signal
is illustrated in Figure 3. The signal had a smooth,
sinusoidal pattern that varied throughout the model heart
cycle. The signal was broadly similar, but had
an additional spiked-wave pattern that coincided with
the simulated ventricular depolarisation. The difference
between the two ventricular volumes measurements,
ΔVECG during this phase of the cardiac cycle varied
between –12% (i.e. an underestimation) and +9%. By
contrast, the difference during the remainder of the car-
diac cycle varied only slightly from –0.3 to +0.9%.
The pressure-volume loop obtained with the sig-
nal had a quadrilateral shape with four distinct phases (Fig-
ure 4). End-diastole and end-systole were each identifiable
as the single pressure-volume point at the lower right-hand
and upper left-hand corners, respectively. The ECG in-
terference pattern altered the shape of the pressure-volume
loop, primarily affecting the late filling and early iso-
volumic contraction phases. As a result, end-diastole
Copyright © 2009 JBiSE
S. P. McGuirk. Lu et al. / J. Biomedical Science and Engineering 2 (2009) 491-498
Figure 3. Calibrated conductance volume measurements
versus time. Conductance volume measurements were
made with (; A) and without ECG interference
(; B) versus time. The ECG signal (ECGinput) has
been plotted (C) for comparison.
could not be reliably identified (Figure 4).
Comparable results were obtained under all experi-
mental conditions. Increasing the end-diastolic volume
from 385 ml to 500 ml did not significantly change the
discrepancy between the two ventricular volume meas-
urements. While increasing the resistivity from 37 ·cm
to 330 ·cm increased the median measurement dis-
crepancy slightly, from +0.2% to +1.8%, it did not sig-
nificantly alter the maximal range of the discrepancy.
The conductance catheter technique is an established
method that enables continuous volume measurements
based on the electrical conductance of the intraventricular
blood pool. This study has demonstrated that other electri-
cal signals within the ventricular cavity alter the measured
conductance. This produces a conductance signal “arte-
fact”, which varies as a function of the first-derivative of
the additional electrical signal. This artefact represents a
novel and additional source of error that potentially af-
fects the accuracy of ventricular volume measurements
made using the conductance catheter technique.
Ventricular depolarisation and repolarisation cause a
measurable electrical signal within the ventricular cavity
[12]. In the present study, a simulated ventricular elec-
trogram produced a biphasic signal with a highamplitude
spiked wave pattern that coincided with the QRS com-
plex and a comparatively low-amplitude sine wave- like
pattern during the T wave. This signal was associated
with a conductance volume measurement error that
ranged between a 12% volume underestimation to a 9%
volume overestimation. The entire range of this meas-
urement discrepancy occurred within a 24 ms period
during simulated ventricular depolarisation. By contrast,
simulated repolarisation was associated with a small,
clinically unimportant measurement error. The precise
pattern, will vary with the morphology of ventricular
electrogram [12,13,14]. The precise ECG interference
pattern, GECG(t) will vary with the morphology of ven-
tricular electrogram [12,13,14]. Nevertheless, these in
Figure 4. Pressure-conductance volume loop from the model
ventricle. Conductance volumemeasurements were made
with (; black line) or without ECG interference (;
grey dashed line).
Copyright © 2009 JBiSE
S. P. McGuirk et al. / J. Biomedical Science and Engineering 2 (2009) 491-498 497
vitro findings are consistent with our previously unre-
ported clinical findings.
The haemodynamic events during the cardiac cycle
are best displayed by plotting the instantaneous left ven-
tricular pressure versus volume [15]. Under steady-state
conditions, this pressure-volume loop has a quadrilateral
shape where each side represents one of four functional
distinct phases: filling, isovolumic contraction, ejection
and isovolumic relaxation. End-diastole and end-systole
are identifiable as the single pressure-volume points in
the lower right-hand and upper left-hand corners, re-
spectively. However, ventricular depolarisation
the rapid rise in intraventricular pressure that marks the
onset of ventricular systole. The conductance signal ar-
tefact identified in this study meant that end-diastole
could no longer be reliably identified on the pressure-
conductance volume loop alone.
End-diastole may alternatively be defined using the
surface electrocardiogram as the onset of the QRS com-
plex [16]; the R wave peak [17]; or up to 40 ms after the
R wave peak [18]. End-diastole may also be defined as
the R wave peak on the ventricular electrogram. In our
experience, this time-point occurs synchronously with
the onset of ventricular systole [19]. However, conduc-
tance volume measurements at all of these time-points
will be variably affected by the conductance signal arte-
fact such that end-diastolic volume cannot be accurately
measured using the conductance catheter technique. This
in turn means that indices of ventricular function that are
based on EDV, such as cardiac output, ejection fraction
together with the quantitative assessment of ventricular
compliance, will be adversely affected as a consequence
of the conductance signal artefact.
The limitations of the physical model have been de-
scribed previously [10]. Electrical activity within the
ventricle was represented using a fixed dipole within the
ventricular cavity. This comparatively simple model en-
abled characterisation and quantification of a new con-
ductance measurement error. However, the model did
not include any representation of the ventricular wall and
the effect of parallel conductance was not examined. A
moving dipole or multiple dipoles within an artificial
ventricular wall would also have provided a more
physiological model.
This study has demonstrated that the accuracy of these
conductance volume measurements is adversely affected
by other electrical signals, such as the ventricular elec-
trogram. The ventricular electrogram produced a clini-
cally important volume measurement that meant end-
diastole could neither be precisely identified nor accu-
rately measured. These original findings have important
implications for the quantitative assessment of ventricu-
lar function and, in particular the assessment of chamber
Simon McGuirk was supported by a British Heart Foundation Junior
Research Fellowship (FS/03/102).
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