IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 54, NO. 4, APRIL 2007

Capacitive Sensing of Electrocardiographic Potential
Through Cloth From the Dorsal Surface of the Body in a
Supine Position: A Preliminary Study
Akinori Ueno*, Yasunao Akabane, Tsuyoshi Kato,
Hiroshi Hoshino, Sachiyo Kataoka, and Yoji Ishiyama

Abstract—A method for obtaining electrocardiographic potential
through thin cloth inserted between the measuring electrodes and theskin of a subject’s dorsal surface when lying supine has been proposed.
The method is based on capacitive coupling involving the electrode, thecloth, and the skin. Examination of a pilot device which employed themethod revealed the following: 1) In spite of the gain attenuation in thehigh frequency region, the proposed method was considered useful formonitoring electrogardiogram (ECG) for nondiagnostic purpose. 2) The
method was able to yield a stable ECG from a subject at rest for at least 7h, and there was no significant adverse effect of long-term measurementon the quality of the signal obtained. 3) Electrode area was the factor that
had most influence on the signal, compared with other factors such as
cloth thickness and coupling pressure, but could be reduced to 10 cm
for heart rate detection. 4) Input capacitance of the device was assumed
to be the dominant factor for the gain attenuation in the high frequency
region, and should be reduced with a view to diagnostic use. Although
there is still room for improvement in terms of practical use, the proposed
method appears promising for application to bedding as a noninvasive andawareness-free method for ECG monitoring.
Index Terms—Home health care, insulator electrode, nonobtrusive
monitoring, unconstrained and noninvasive ECG measurement, wearable
sensing.

I. INTRODUCTION
The gradual aging of society in Japan, and the resulting increasein healthcare expenditure, has highlighted the need for efficient systems for monitoring of elderly people in their homes over long periodsof time. Recording of physiological variables, such as the electrocardiogram (ECG), during everyday life could be useful for managementof individuals with chronic health disorders [1]. Furthermore, real-lifelong-term health monitoring could be helpful for assessing the effectsof treatment at home, and would also be potentially beneficial for observing deviations in health status from the norm at an early stage, or for
automatically alerting paramedics in emergency cases. Consequently,
long-term home health monitoring could be valuable not only for improving the quality of life (QOL) of senior citizens but also for reducing
health care expenditure.

Manuscript received March 14, 2006; revised August 29, 2006. This work
was supported in part by “Academic Frontier” Project for Private Universities:
matching fund subsidy from the Ministry of Education, Culture, Sports, Scienceand Technology (MEXT), 2003–2004, and in part by the New Energy and Industrial Technology Development Organization (NEDO) of Japan through theIndustrial Technology Research Grant Program in 2005–2006. Asterisk indicates corresponding author.
*A. Ueno is with the Department of Electronic and Computer Engineering,Tokyo Denki University, Ishizaka, Hatoyama-machi, Saitama 350-0394, Japan(e-mail: ueno@f.dendai.ac.jp).
Y. Akabane and T. Kato are with the Master’s Program in Electronic andComputer Engineering, Tokyo Denki University, Saitama 350-0394, Japan.
H. Hoshino is with the Department of Electronic and Computer Engineering,Tokyo Denki University, Saitama 350-0394, Japan.
S. Kataoka is with the Apprica Childcare Institute, Tokyo 104-0061, Japan.
Y. Ishiyama is with the School of Health Sciences, Kyorin University, Tokyo
192-8508, Japan.
Digital Object Identifier 10.1109/TBME.2006.889201

759

ECG has been part of routine cardiovascular evaluation for almost
three decades, and is now recognized to have potential for long-termmonitoring at home over periods of weeks or even months. Since conventional systems for ECG monitoring are unsuitable for long-term
studies, previous researchers have focused on developing dedicatedsystems for long-term monitoring in daily life [2]–[11]. For example,
Kwatra [2], [3], Ishijima [4], and Tamura [5] have developed systems
for ECG or heart rate monitoring during bathing. Ishijima [6], [7] developed an automated system for ECG monitoring during sleep using
embedded textile electrodes in a bed set. Park [8] proposed a prototype
intelligent garment that can collect, process, store, and transmit information about the wearer, such as ECG parameters, and Catrysse [9],Scilingo [10], and Paradiso [11] have implemented this idea by using
textile or fabric sensors. Since a noninvasive and awareness-free approach is essential for long-term home monitoring, the idea of sensors
embedded in furnishings or daily clothing is quite rational. However,there are still some challenges to be addressed regarding the utility ofsuch systems.
In conventional ECG measurement, an electrolytic paste or a conductive adhesive is almost always required for maintaining reliableohmic contact with the skin. Therefore, ECG measurement for a long
period using conventional methods causes irritation and discomfort,
and is a potential cause of skin allergy and inflammation. Despite thesedisadvantages of conventional ECG detection, sufficient consideration
has not been given to skin-to-electrode coupling [8]. Previous authors[6], [7], [9]–[11] have addressed this problem by employing textile
electrodes, which do not require any electrolytic paste or conductive
adhesive for measurement, to relieve potential irritation and discomfort. Nevertheless, the potential for metal allergy or discomfort upon
direct long-term contact of the skin with metal material has not yet
been addressed.
With the above background in mind, the authors have focused on theprinciple of capacitive sensing capable of detecting alternating electrical potential through an inserted thin insulator, and have applied theprinciple for measurement of electrocardiographic potential throughcommonly available cloth from the dorsal surface of the body when
the subject is supine. In order to explore the feasibility and applicability of this idea, the authors have fabricated a pilot measuring device
and conducted some preliminary experiments and model studies usingit.
II. CAPACITIVE SENSING
Capacitive sensing is based on the principle of the capacitive (or insulated) electrode [12]–[18], which is an electrode not requiring the useof electrolytic paste or conductive adhesive. The electrode relies on capacitive coupling, conventionally involving a metal electrode, an insulator, and the skin, as shown in Fig. 1(a). The electrode can carry an alternating bioelectric current equivalently through the capacitance of thecoupling. In previous studies, various insulators, such as anodized aluminum [12], [13], silicon dioxide [14], [15], pyre varnish [16], anodic
insulated tantalum oxide [17], or barium titanate [18], have been tested
as these materials exhibit high permittivity. In the present study, commonly available cloth, especially cotton, was substituted for these rigid
insulators to relieve the irritation, allergy and discomfort experienced
with conventional skin-to-electrode coupling, as shown in Fig. 1(b).
Also a sheet of conductive fabric was substituted for the conventional
metal electrode so as to achieve a deformable coupling corresponding
to the contour of the coupled region.
Capacitive sensing is also based on an impedance matching circuit to
mediate the high impedance of the coupling with a low impedance required by the subsequent circuitry. For example, a field-effect transistor

0018-9294/$25.00 © 2007 IEEE

760

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 54, NO. 4, APRIL 2007

Fig. 1. Schematic models of (a) conventional and (b) modified capacitive electrodes coupled to the skin. " and " indicate dielectric constants of the inserted
insulator and the cloth, respectively. " is the permittivity of free space.

Fig. 2. Configuration of measuring electrodes and block diagram of the present system.

(FET) source follower has been used as an impedance matching circuit
and actively mounted in an electrode housing [15], [16] for reducing
common mode interference caused by power lines. In the present study,
an instrumentation amplifier [19] composed of operational amplifierswith high input impedance (National Semiconductor, LF356, 1000 G
according to the specification sheet)1 was employed in order to matchthe greater impedance of the coupling due to the smaller dielectric
constant of the inserted cloth in comparison with conventional insulators. The impedance matching circuit was not mounted in the electrodehousing to avoid hard contact with the skin.
III. MATERIALS AND METHODS
A. Fabric ElectrodeThe electrocardiographic signal is picked up by electrodes made of asquare sheet of conductive fabric with conductive acrylic adhesive (3M,2191FR). The electrodes are stuck to a mattress with the adhesive andcovered by a cotton bed sheet. Two lead electrodes are located under theright and left scapulae of the subject when lying in a supine position,
as shown in Fig. 2. A reference electrode is placed beneath the rightlumbar region. Skin-cloth-electrode coupling is held by the subject’s
weight on the cloth and by repulsive force from the mattress. The leadelectrodes are connected to a measuring device, as described in the next
section, by using shielded wires 2.0 m long. Three sets of measuring
electrodes, each with an area of 70, 40, or 10 cm2 , were prepared for
examining the influence of the electrode area on an obtained signal.
B. Signal Extraction
The pilot measuring device with filtering and amplification circuitry
was fabricated using off-the-shelf components. The device consists ofan instrumentation amplifier, a high-pass filter, two notch filters, a lowpass filter and an inverting amplifier. A block diagram of the device is1[Online]. Available at http://www.national.com/pf/LF/LF356.html.

shown in Fig. 2. The instrumentation amplifier is employed not only
as a differential amplifier but also as the impedance matching circuit.
The circuit elements of the high-pass filter and the low-pass filter are
determined in order to obtain a cutoff frequency of 0.05 and 100 Hz,
respectively. A Butterworth filter (NF, CF-2BE-HB-50Hz-Q5) is used
for the former notch filter in order to reduce 50-Hz interference. The
device is powered by four series D batteries to obviate the possibilityof electric shock. A dc-dc converter (Datel, BST-12/105-d5)2 is used
to boost the nominal 6.0 V supply to a regulated 612 V.
C. Simultaneous Measurement With a Conventional Device
One male subject, aged 22, naked from the waist up, was instructed
to lie supine on the mattress bearing the electrodes with an area of 70
cm2 . A cotton bed sheet 395 m thick was used to cover the electrodes and the mattress. Electrocardiographic potential was recorded
using the developed device from the dorsum of the subject through theinserted cotton sheet. The output signal from the device was digitizedat 1 kHz by an analog-to-digital converter and stored in a personal
computer using a data acquisition system (Biopac Systems, MP-150
system). As a reference signal, the ECG was recorded using a conventional bioamplifier (Teac Instruments, BA1104-CC) and disposable
electrodes (Nihon Kohden, Vitrode F-150M).3 The disposable electrodes were attached directly to the dorsal surface of the subject at
positions close to each electrode on the mattress. As another reference, the Lead I ECG was measured using a standard limb lead withthe same conventional bioamplifier. Both reference signals were wirelessly transmitted and received with a telemeter unit (Teac Instruments,
TU-4), and then digitized simultaneously with the output signal of thedeveloped device.
2[Online]. Available at http://www.datel.com/data/power/bst-3w.pdf.
3[Online]. Available at http://www.nihonkohden.com/products/type/supplies/dispo\_electrodes.html.

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 54, NO. 4, APRIL 2007

761

D. Experimental Long-Term Monitoring Using the Device
In order to examine a stability of the capacitive coupling of our device, an experiment to examine long-term monitoring was conductedusing the same volunteer. The subject, naked from the waist up, was instructed to lie supine at rest for 7 h on the mattress, which was coveredby the same cloth, in the same manner as that described in Section III-C.
An initial measurement was conducted for 15 min, and then every hour
thereafter.
E. Influence of Coupling Condition on the Obtained Signal
If the coupling between an electrode and the skin can be modeled
as in Fig. 1(b), the capacitance C of a coupling having a distance of dmeters, a coupling area of S square meters, and a coupling dielectric
constant " is given by

S
C=" :
d

(1)

Since any change in the capacitance of the coupling is related to achange in impedance (i.e. voltage loss) at the coupling, the output
signal of the developed device is considered to be affected by the areaof the coupling (i.e. electrode area), the distance between the electrodeand the skin, and the dielectric constant of the coupling. Then, theinfluence of coupling condition on the output signal can be examinedby changing the electrode area, the thickness of the inserted cloth, andthe pressure on the coupling.
1) Influence of Electrode Area: Measuring electrodes having an
identical area of 70, 40, or 10 cm2 were placed on the mattress according to the configuration shown in Fig. 2. The electrodes were covered with the cotton sheet 395 m thick. Then, three synthetic human
dorsal surfaces made of square pieces of conductive fabric having thesame area of the measuring electrode were attached to the cloth at locations just above each of the measuring electrodes, so that the whole areaof each measuring electrode was involved in capacitive coupling witheach synthetic dorsal surface. The electrodes, the sheet and the synthetic dorsal surfaces were subjected to 1751 Pa pressure using a weight
to mimic the pressure resulting from a supine adult subject 1.65 m inheight and weighing 60 kg (dorsal area 0.336 m2 ). The simulated ECG
signal was input from an ECG generator (Nihon Kohden, AX-201D)
to the synthetic dorsal surfaces. The output signal from the device, andalso the frequency-gain response from 0.01 to 500 Hz, was measured
for each of the different electrode areas.
2) Influence of Thickness of the Inserted Cloth: Cotton sheets withdifferent thicknesses of 395, 463, or 1020 m at atmospheric pressure
were inserted between the electrode and the skin, and the output signal
from the device was measured for each thickness. Three 70-cm2 electrodes, three synthetic dorsal surfaces, the ECG generator and a pressure of 1751 Pa were used in a similar way to that in Section III-E1).
Also, the frequency responses were measured respectively for each
sheet thickness.
3) Influence of Pressure on the Coupling: The simulated ECG and
the frequency response for each thickness of cotton sheet were measured at a pressure of 1751, 1083, or 578 Pa. The 70-cm2 electrodes,
the synthetic dorsal surfaces and the ECG generator were used in a similar way to that described in Section III-E1). The pressure of 578 Pa
mimics experimentally the conditions created by a 2.36-kg neonate
with a contact area of 0.040 m2 . Additionally, the least pressure capableof detecting the R-wave was investigated for each cloth thickness. The
room temperature was 19  C  22  C and the relative humidity was
46%56%.
4) Influence of Electrode Conductivity: In order to investigate
whether the conductivity of the electrode influenced the results, thesimulated ECG and the frequency response without any inserted cloths
were measured at each pressure. The 70-cm2 electrodes, the synthetic

Fig. 3. Equivalent circuit of the capacitively coupled cloth. R is the contact
resistance between the cloth and the conductive surfaces such as electrode orskin. R and C are dc resistance and capacitance of the cloth, respectively.

dorsal surfaces and the ECG generator were used in a similar way tothat described in Section III-E1).
F. Frequency-Impedance Response of the Inserted Cloth
In order to investigate the interaction between the electricalproperties of the inserted cloth and the obtained signal, the frequency-impedance response was measured for each cloth with a
precision LCR meter (Agilent Technologies, 4284A). The cloth was
fixed at a constant pressure by a dielectric test fixture (Agilent Technologies, 16451B). The frequency was varied from 20 Hz, which is thelower limit of the instrument, to 300 Hz. To complement the deficit of
the raw data below 20 Hz, we used the model of capacitively coupledcloth shown in Fig. 3, and then estimated the model parameters foreach cloth by fitting the following equation to the measured impedanceof the cloth, Zcloth :jZcloth (f )j = Rcn +

Rcl

(2)

1 + (2fCcl Rcl )2

where Ccl and Rcl are the capacitance and dc resistance of the cloth,
Rcn is the contact resistance between the cloth and conductive surfaces,
and f is the frequency.
G. Input Capacitance of the Developed Device
We expected that the input capacitance of the device would eventually degrade the performance in the high frequency region. Since
the input capacitance was too low to be measured with conventional
equipment, we devised a method for assessing it. This procedure was
performed as follows. First, two off-the-shelf capacitors with the same
capacitance were connected at the front ends of the differential inputof the device instead of the two lead electrodes, and the frequency response while the capacitors were inserted was measured. Then, thegain in decibels at each frequency with the capacitors inserted was subtracted from the corresponding original gain measured without the capacitors. Second, the developed device connected to the two capacitors
was modeled by an equivalent circuit, as shown in Fig. 4. In the figure,the gain of the device with the capacitors, Gcap (f ) = vout =vin , can be
0 ,expressed using the original gain of the device, Gorg (f ) = vout =vin
and circuit components in Fig. 4, as follows:

Gcap (f ) =j 2fRin C0
Gorg (f )2 + j 2fRin (2Cin + C0 )(3)where C0 is the capacitance of the externally inserted capacitors, andCin and Rin are the input capacitance and input resistance of the device,respectively. According to (3), the gain difference in decibels between
20 log jGcap j and 20 log jGorg j, i.e. 20 log(jGcap j=jGorg j), can be described as follows:
20 logjGcap (f )j = 20 log
jGorg (f )j2fRin C0

2 (2C + C )2
4 + 4 2 f 2 Rin
0
in

:

(4)

762

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 54, NO. 4, APRIL 2007

Fig. 4. Equivalent circuit of the input part of the developed device in which
capacitors are inserted at the front ends. C is the capacitance of the inserted
capacitor. C and R are the input capacitance and input resistance of thedevice, respectively.

Fig. 6. Long-term recordings using the present device from the dorsum withan inserted cotton sheet 395 m thick. S/N ratio of the recordings after 4 h was
as good or better than that at the 4-h point. The room temperature was 21 C
and the relative humidity was 24%.

Fig. 5. Recordings obtained simultaneously using the present device witha cotton sheet 395 m thick from the dorsum (top), using a conventional
bioamplifier without a cloth from the dorsum (middle), and using a conventional bioamplifier and a standard limb lead without a cloth (bottom). The top
recording was subjected to a 20-point moving average to eliminate power line
noise.

Since the value of C0 is known and Rin can be obtained from the specification sheet of the IC used in the device, we can consider (4) as afunction of f and Cin . Therefore, we can estimate Cin by fitting (4) tothe experimentally obtained gain difference between 20 log jGcap j and
20 log jGorg j.
In the actual experiment, ceramic 102-pF or 478-pF capacitors wereconnected individually via the lead wires, and the two frequency responses were measured. Moreover, frequency responses not including
the wires were also measured. The capacitance of the capacitors wasmeasured with the precision LCR meter.
IV. RESULTS
A. Simultaneous Measurement With a Conventional Device
Fig. 5 shows recordings typical of those obtained. The output signalof the developed device showed periodic waveforms synchronized to,and similar to, the reference ECGs. Therefore, the proposed method
was considered useful at least for detecting heart rate and/or other periodic parameters of the ECG even with the cloth inserted. However, theratios of the peak amplitude of the P-wave and T-wave to the R-wave
in the top recording appeared to be slightly greater than those in thereference ECGs. This distortion was probably due to inconstant gain
through the pass band of our device when the cloth was inserted.
B. Experimental Long-Term Monitoring Using the Device
As can be seen in the top recording in Fig. 6, depressed R-waves
and skewed T-waves were observed at the beginning of the experiment.

Fig. 7. Typical recordings contaminated by motion artifacts during long-term
measurement. (a) Recording made at the beginning, and (b) 2 h later. The top
recording in each figure is that obtained using the present device from thedorsum with a cotton sheet 395 m thick inserted, and the bottom recordingis the wirelessly measured lead I ECG.

However, they had disappeared before measurement was conducted at
the 1-h point. In addition, there was a marked decrease in hum noise inrecordings after 4 h, approaching a level as low as that in the directly
measured ECG.
Fig. 7 shows typical recordings contaminated by motion artifacts atthe beginning, and 2 h later. Since the impedance of the coupling is assumed to have been high at the beginning, as can be seen from the low
signal-to-noise (S/N) waveform in Fig. 6, the output signal from thedevice at the beginning was susceptible to motion artifacts. As a result,
it is evident at the beginning of the experiment that the motion of thesubject contaminated the output from the developed device, but not thelead I ECG from the conventional electrocardiograph [see Fig. 7(a)].
However, once the initial impedance of the coupling had decreased
after a while, the output of the device seemed to be more stable thana conventional electrocardiograph. As seen in Fig. 7(b), limb movements contaminated the lead I ECG, but not the output signal from thedeveloped device, contrary to the situation observed at the beginning.

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 54, NO. 4, APRIL 2007

763

Fig. 8. Influence of (a) electrode area, (b) cloth thickness, and (c) coupling pressure on ECG recording (left column) and on frequency response (right column).
The frequency response labeled “original” corresponds to that obtained directly from the measuring electrode. The gain at 50 Hz was sharply attenuated because
50-Hz band elimination filters were used in the system. The original frequency response satisfied the Japanese Industrial Standard for an electrocardiograph (JIST1201), except around the 50-Hz frequency band.

This phenomenon was often observed in recordings after 2 h. Therefore, there was no significant adverse effect of long-term measurementon the quality of the output signal from the developed device.
C. Influence of Coupling Condition on the Output Signal
1) Influence of Electrode Area: As shown in Fig. 8(a), the periodic R-wave was evident in all recordings, even for the smallest electrode area of 10 cm2 . However, the waveform of the PQRST com-plex became distorted as the electrode area decreased. Frequency responses in Fig. 8(a) suggests that this distortion was attributable toa tilted frequency-gain response in the 0.1–100 Hz region. Since thegain in the high-frequency domain decreased as the area of the electrode decreased, the R-wave was depressed as the area decreased andthe S-wave was absent in the recording obtained with the 10-cm2 electrodes. Therefore, the area can be reduced to 10 cm2 if only heart rate
is being detected.

764IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 54, NO. 4, APRIL 2007

TABLE IESTIMATED VALUES OF CONTACT RESISTANCE (R ), dc RESISTANCE (R ), CAPACITANCE (C ), AND RELATIVEPERMITTIVITY OF THE INSERTED CLOTHS. RELATIVE PERMITTIVITY WAS ESTIMATED BY USING (1)Fig. 9. Frequency-impedance responses of the inserted cloths. Solid lines in the
figure are the best fitted curves obtained by applying (2). Estimated parameters
are shown in Table I. The room temperature during the measurement was 24 C
and the relative humidity was 40% 46%.

2) Influence of Thickness of the Inserted Cloth: As can be seen in
Fig. 8(b), the PQRST complex was sufficiently recognizable even witha cloth thickness of 1020  , and the quality was similar to that of otherrecordings obtained with thinner cloth. Although the gain had a tendency
to decrease as the frequency increased, all the frequency responses obtained with the cloth indicated a close gain for each frequency. In fact the
responses obtained with the 463  cloth and the 1020  cloth were
almost the same. This may be because the coupling impedance of thesecloths is almost the same at a pressure of 1751 Pa. On the other hand, theleast gain was obtained for the 395- cloth, despite the fact that thiswas the thinnest. This minimal gain for the 395  cloth was presumably due to the low dielectric constant and the high resistivity of the cloth.
Thus, it appears that the thickness of the inserted cloth at atmosphericpressure can be increased to at least 1000  as long as the electrode iscoupled firmly with sufficient pressure.
3) Influence of Pressure on the Coupling: As indicated in Fig. 8(c),reduction of the pressure from 1751 to 578 Pa had little influence on
both the S/N and the waveform of the recordings for the 395  cloth.
The frequency responses, shown in Fig. 8(c), agreed with the results.
With regard to the other two cloths, little effect was also confirmed.
Moreover, the R-wave could be recognized for each cloth thickness
even when no weight was placed on the electrodes. Thus, the developed
device shows promise for application not only to adult subjects, but also
to neonates for heart rate monitoring.
4) Influence of Electrode Conductivity: No gain attenuation of theelectrodes was confirmed from 0.01 to 500 Hz. Therefore, the gain attenuation observed in Fig. 8 was not due to gain loss at the electrodebut to that at the coupling involving the cloth.

m

m

m

m

m

m

m

Fig. 10. Frequency response of (a) the gain with capacitors inserted at the front
ends of the system and (b) the difference in gain between (a) and the original
response. The original frequency response is that measured directly from theelectrodes. The solid and dotted lines in (b) are the best fitted curves obtainedby applying (4).

E. Input Capacitance of the Developed Device
As shown in Fig. 10, the plots obtained when the lead wires were
included were apparently smaller than those excluding the wires. As aconsequence, Cin of 370 and 65 pF indicated the least root mean squareerror for the results including and excluding the wires, respectively.
Thus, the lead wires had an input capacitance of 305 pF.
V. DISCUSSION
A. Perspiration of the Subject

D. Frequency-Impedance Response of the Inserted Cloth
In accordance with the results shown in Fig. 9, impedance of the395  cloth was estimated to be the highest in a frequency region
below 10 Hz. Table I indicates this was due to the largest dc resistanceand the least relative permittivity of the cloth. Thus, the least gain forthe 395  cloth in Fig. 8(b) can be attributed to these electrical properties of the cloth.

mm

In a previous article, Geddes [20] reported that “dry” electrodes arereally only dry when first applied because of subsequent perspiration,and that impedance at 20 min was between one-fourth and one-fifth of
the initial impedance. Therefore, the improvement of the S/N with timein Fig. 6 was probably due to a decrease in both the impedance of thecloth and the contact resistance between the cloth and the skin, originating from the perspiration of the subject. In long-term measurement,

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 54, NO. 4, APRIL 2007765

device with cloth inserted at the front ends can be modeled by substituting C0 in Fig. 4 for Zcloth in Fig. 3, the gain with the cloth inserted,
Gcloth (f ), can be described as follows:

Gcloth (f )(j 4fCin + 2=Rin )01 Gorg (f )= R + (j 2fC
01 + (j 4fCin + 2=Rin )01 :cn
cl + 1=Rcl )1(5)

Fig. 11. Frequency responses of the simulated gain using (5) with the 395-m
cloth inserted. The original frequency response was that measured directly fromthe electrodes. Parameter values for estimating the response are as follows:
R = 312 k , C = 365 pF, R = 34 M , C = 65 or 370 pF,
R = 1 T . R , R , and C were compensated from the values in Table I
by considering the electrode area of the precision LCR meter (11.3 cm ) andof the fabric lead electrodes (70 cm ).

We can estimate the frequency-gain responses with the cloth inserted
by substituting the parameter values in (5) and by using the experimentally obtained Gorg (f ). As shown in Fig. 11, the simulated responsewhen Cin equals 370 pF exhibited a characteristic similar to the experimentally obtained responses shown in Fig. 8. Moreover, the simulation when Cin equals 65 pF predicted that gain attenuation in thehigh-frequency regions could be suppressed by decreasing the input
capacitance. Therefore, it is necessary to reduce the input capacitance,
for example by shortening the wire length between the electrodes andthe impedance matching circuits, in order to acquire a less distorted
waveform and to make the device applicable for diagnostic use.
VI. CONCLUSION

once a perspiration layer has formed, there will be a physical contact
between the electrodes and the skin. Therefore, for the next step, thecurrent nickel-plated conductive fabric should be substituted by anotherfabric containing nonallergenic conductive materials.

B. Motion Artifact During Measurement
As shown in Fig. 7(b), limb movements contaminated the lead
I ECG, but not the output signal from the developed device, contrary
to the situation observed at the beginning. We consider that threepossible factors can explain this phenomenon. The first factor is thegreater input impedance of the present device. The input impedance
was estimated to be greater than 430 M in a frequency region
below 1 Hz, even with an input capacitance of 370-pF, by using theequivalent circuit in Fig. 4. Thus, once resistive coupling had beenaccomplished at the electrode-to-skin coupling by perspiration, theslow sub-1-Hz component of the change in contact impedance caused
by limb motion would have had less impact on the obtained signal
than if a conventional device had been used. The second factor isthat the lead wires of the developed device were connected not to theelectrodes on the peripheral limbs but to those on the mattress under
the trunk. Accordingly, subtle limb motion would hardly disturb thelead wires of the developed device. The third factor is detachment of
the sides of the disposable electrode pad due to perspiration, perhaps
reducing the adhesion of the electrode.
Thus, future issues that need to be addressed are shifting of the trunk
off the electrodes by larger motions such as rolling over, and the motion artifact at the beginning of measurement. One way of coping withthese issues would be introduction of multiple arrays of horizontal strip
electrodes, and of filters and software to search for stable electrodesets, similar to the approach reported by Niizeki [21]. Another method
would be to narrow the filter band width for heart rate monitoring.

C. Gain Attenuation in the High Frequency Region
As shown in Fig. 8, the frequency responses with the cloth inserted
indicated that the gain of the device decreased as the frequency increased. This attenuation is perhaps due to the input capacitance of thedeveloped device. Since an equivalent circuit of the input part of theWe have proposed an approach for obtaining electrocardiographic
potential from the dorsal surface of a subject lying supine under conditions whereby a thin cloth is inserted between the electrodes and theskin. We fabricated a pilot measuring device and explored the feasibility and applicable scope of the proposed method. Examination of
the device yielded the following results.
1) In spite of the gain attenuation in the high frequency region, theproposed method was considered useful for monitoring ECG for
nondiagnostic purposes.
2) The method was able to yield a stable ECG from a subject at rest
for at least 7 h, and there was no significant adverse effect of longterm measurement on the quality of the signal obtained.
3) Electrode area had a greater influence on the signal than cloth
thickness and coupling pressure, but could be reduced to 10 cm2
for heart rate detection.
4) Input capacitance of the device was assumed to be the dominantfactor for the gain attenuation in the high frequency region, andthis should be reduced to make the device applicable for diagnostic use.
Although there is still room for improvement in terms of its practicaluse, the proposed method appears promising for application to bedding
as a noninvasive and awareness-free method for ECG monitoring.
ACKNOWLEDGMENT
The authors would like to thank to Y. Shiogai in our laboratory forassistance with some of the fundamental experiments, and also theanonymous reviewers for valuable comments and suggestions that improved the manuscript.

REFERENCES[1] I. Korhonen, J. Parkka, and M. V. Gils, “Health monitoring in the homeof the future,” IEEE Eng. Med. Biol. Mag., vol. 22, pp. 66–73, May/Jun.
2003.
[2] S. C. Kwatra and V. K. Jain, “A new technique for monitoring heartsignals—part I: instrumentation design,” IEEE Trans. Biomed. Eng.,
vol. BME-33, no. 1, pp. 35–41, Jan. 1986.
[3] ——, “A new technique for monitoring heart signals—part II: orthogonal lead extraction,” IEEE Trans. Biomed. Eng., vol. BME-33, no. 1,
pp. 1–9, Jan. 1986.
[4] M. Ishijima and T. Togawa, “Observation of electrocardiograms
through tap water,” Clin. Phys. Physiol. Meas., vol. 10, no. 2, pp.
171–175, 1989.

766

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 54, NO. 4, APRIL 2007

[5] T. Tamura, T. Yoshimura, K. Nakajima, H. Miike, and T. Togawa, “Unconstrained heart-rate monitoring during bathing,” Biomed. Instrum.
Technol., vol. 31, no. 4, pp. 391–396, Jul. 1997.
[6] M. Ishijima, “Monitoring of electrocardiograms in bed without utilizing body surface electrodes,” IEEE Trans. Biomed. Eng., vol. 40,
no. 6, pp. 593–594, Jun. 1993.
[7] ——, “Cardiopulmonary monitoring by textile electrode without subject-awareness of being monitored,” Med. Biol. Eng. Comput., vol. 35,
pp. 685–690, Nov. 1997.
[8] S. Park and S. Jayaraman, “Enhancing the quality of life through wearable technology,” IEEE Eng. Med. Biol. Mag., vol. 22, pp. 41–48, May/
Jun. 2003.
[9] M. Catrysse, R. Puers, C. Hertleer, L. Van Langenhove, H. van
Egmondc, and D. Matthys, “Towards the integration of textile sensors in a wireless monitoring suit,” Sens. Actuators A, vol. 114, pp.
302–311, Jan. 2004.
[10] E. P. Scilingo, A. Gemignani, R. Paradiso, N. Taccini, B. Ghelarducci,
and D. D. Rossi, “Performance evaluation of sensing fabrics for monitoring physiological and biomechanical variables,” IEEE Trans. Inf.
Techn. Biomed., vol. 9, no. 3, pp. 345–352, Sep. 2005.
[11] R. Paradiso, G. Loriga, and N. Taccini, “A wearable health care system
based on knitted integrated sensors,” IEEE Trans. Inf. Techn. Biomed.,
vol. 9, no. 3, pp. 337–344, Sep. 2005.
[12] P. C. Richardson, F. K. Coombs, and R. M. Adams, “Some new electrode techniques for long term physiologic monitoring,” Aerosp. Med.,
vol. 39, pp. 745–750, Jul. 1968.
[13] A. Lopez, Jr. and P. C. Richardson, “Capacitive electrocardiographic
and bioelectric electrodes,” IEEE Trans. Biomed. Eng., vol. BME-16,
no. 1, p. 99, Jan. 1969.
insulated[14] R. N. Wolfson and M. R. Neuman, “Miniatureelectrodes based on semiconductor technology,” presented at the 22nd
ACEMB, Chicago, Ill., 1969.
[15] W. H. Ko, M. R. Neuman, R. F. Wolfson, and E. T. Yon, “Insulated
active electrode,” in Proc. Int. Conf. IEEE Solid-State Circuits, 1971,
pp. 195–198.
[16] A. Potter and L. Menke, “Capacitive type of biomedical electrodes,”IEEE Trans. Biomed. Eng., vol. BME-17, pp. 350–351, Oct. 1970.
[17] C. H. Lagow, K. J. Sladek, and P. C. Richardson, “Anodic insulated
tantalum oxide electrocardiograph electrodes,” IEEE Trans. Biomed.
Eng., vol. BME-18, pp. 162–164, Mar. 1971.
[18] T. Matsuo, M. Esashi, and K. Iinuma, “Capacitive electrode forbiomedical use—the use of Barium-titanate ceramics for biomedicalsensing electrode-,” (in Japanese) Jpn. J. Med. Electron. Biol. Eng.,
vol. 11, no. 3, pp. 10–16, Jun. 1973.
[19] R. B. Northrop, Introduction to Instrumentation and Measurements,
2nd ed. Boca Raton, FL: CRC Press, 2005, pp. 68–71.
[20] L. A. Geddes and M. E. Valentinuzzi, “Temporal changes in electrode
impedance whie recording the electrocardiogram with dry electrode,”Ann. Biomed. Eng., vol. 1, pp. 356–367, 1973.
[21] K. Niizeki, I. Nishidate, K. Uchida, and M. Kuwahara, “Unconstrained
cardiorespiratory and body movement monitoring system for homecare,” Med. Biol. Eng. Comput., vol. 43, pp. 716–724, 2005.

Wavelet Packets Feasibility Study for the
Design of an ECG Compressor
Manuel Blanco-Velasco*, Fernando Cruz-Roldán,Juan Ignacio Godino-Llorente, and Kenneth E. Barner

Abstract—Most of the recent electrocardiogram (ECG) compression
approaches developed with the wavelet transform are implemented using
the discrete wavelet transform. Conversely, wavelet packets (WP) are notextensively used, although they are an adaptive decomposition for representing signals. In this paper, we present a thresholding-based method to
encode ECG signals using WP. The design of the compressor has been
carried out according to two main goals: 1) The scheme should be simple
to allow real-time implementation; 2) quality, i.e., the reconstructed signal
should be as similar as possible to the original signal. The proposed scheme
is versatile as far as neither QRS detection nor a priori signal informationis required. As such, it can thus be applied to any ECG. Results show that
WP perform efficiently and can now be considered as an alternative in
ECG compression applications.
Index Terms—ECG compression, electrocardiogram, filter bank, subband coding, thresholding, wavelet coding, wavelet packets (WP), wavelet
transform.

SiSiO

I. INTRODUCTION
The electrocardiogram (ECG) is widely used because it is a noninvasive way to establish clinical diagnosis of heart diseases. Therefore,
ECG processing has been a topic of great interest and is most commonly used in applications such as monitoring. Long-term records havealso become commonly used to extract or detect important informationfrom the heart signals. In these cases, the quantity of data grows significantly and compression is required for reducing the storage required
and transmission times.
A large amount of lossy ECG compression methods have beendeveloped using many different techniques. In [1], a classificationof compression schemes into three categories is proposed: Direct
methods, transform methods, and other compression methods. Mostof the methods classified in the second group are based on the wavelet
transform (WT) and multiresolution analysis, enabling them to obtain
good compression ratios (CRs) and real-time implementation. The set
partitioning in hierarchical trees (SPIHT) algorithm [2] is the mostwidely known algorithm that fulfills these features. Subsequently, it
has been surpassed by other approaches [3], [4] that give high CR,
with the tradeoff being higher computational complexity.

Manuscript received November 28, 2005; revised September 3, 2006. This
work was supported in part by the Ministerio de Educación y Ciencia underGrant PR-2005-0131 and in part by the Fondo de Investigación Sanitaria under
Project PI052277. Asterisk indicates corresponding author.
*M. Blanco-Velasco is with the Department of Teoría de la Señal y Comunicaciones, Universidad de Alcalá, Campus Universitario, 28871 Alcalá deHenares, Madrid, Spain (e-mail: manuel.blanco@uah.es).
F. Cruz-Roldán is with the Department of Teoría de la Señal y Comunicaciones, Universidad de Alcalá, Campus Universitario, 28871 Madrid, Spain(e-mail: fernando.cruz@uah.es).
J. I. Godino-Llorente is with the Department of Ingeniería de Circuitos y
Sistemas, Universidad Politénica de Madrid, 28031 Madrid, Spain (e-mail:igodino@ics.upm.es).
K. E. Barner is with the Department of Electrical and Computer Engineering,
University of Delaware, Newark, DE 19716 USA (e-mail: barner@ee.udel.edu).
Color versions of one or more of the figures in this paper are available online
at http://ieeexplore.ieee.org.
Digital Object Identifier 10.1109/TBME.2006.889176
0018-9294/$25.00 © 2007 IEEE