Sensors and Materials, Vol. 24, No. 6 (2012) 335–346
MYU Tokyo
S & M 0889

Feasibility of Capacitive Sensing of Surface
Electromyographic Potential through Cloth
Akinori Ueno*, Toshihiko Yamaguchi1, Toshiya Iida1,
Yutaka Fukuoka2,**, Yoshinori Uchikawa3 and Makoto Noshiro4
Department of Electrical and Electronic Engineering, Tokyo Denki University,
Kanda-Nishiki-cho, Chiyoda-ku, Tokyo 101-8457, Japan
1Department of Electronic and Computer Engineering, Tokyo Denki University,
Ishizaka, Hatoyama-machi, Saitama 350-0394, Japan
2Graduate School of Biomedical Science, Tokyo Medical and Dental University,
Yushima, Bunkyo-ku, Tokyo 113-8510, Japan
3Division of Electrical and Mechanical Engineering, Tokyo Denki University,
Ishizaka, Hatoyama-machi, Saitama 350-0394, Japan
4Department of Clinical Engineering, School of Allied Health Sciences, Kitasato University,
Kitasato, Sagamihara-shi, Kanagawa 228-8555, Japan(Received July 21, 2011; accepted January 16, 2012)
Key words: capacitive electrode, EMG measurement, guarding technique, drive-ground-plane
technique, human-machine interfaces

A capacitive sensing method was developed for sensing surface electromyographic
potential (sEMG) through a thin cloth. The method is based on capacitive coupling
involving fabric electrodes, the cloth, and the skin of the subject. A proposed
measuring device was assembled by modifying the previously developed instrument forelectrocardiographic potential. Experimental results showed that sEMG obtained withthe proposed system was clearly visible and showed ﬁrings synchronized with sEMG
simultaneously measured with a commercial device and that the spectral powers of bothsEMGs were almost identical. These results demonstrated the potential of the proposed
method to measure sEMG through cloth. The method enables the discomfort due toconventional skin-to-electrode coupling to be eliminated, although the signal quality,
such as the signal-to-noise ratio (S/N), is slightly decreased with the cloth barrier.

1.

Introduction

Since the late 1970s, myoelectric potential (or electromyogram: EMG) has beenstudied and used for biosignal-based mechatronics interfaces such as artiﬁcial limb
Corresponding author: e-mail: ueno@mail.dendai.ac.jp
Present affiliation: Department of Electrical Engineering, Kogakuin University, Tokyo, Japan

*

**

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prostheses,(1–4) robot hands,(5) and manipulators.(6) Recently, the EMG has opened up
new possibilities for other human-computer interfaces in the ﬁelds of augmentative
and alternative communication(7–10) and environmental control.(11) The EMG contains
information on muscle recruitment, timing pattern, and force. Therefore, the EMG isuseful as an input signal for human adaptive mechatronics.(12)In the conventional surface EMG (sEMG) measurement, an electrolytic paste or aconductive adhesive is almost always required for maintaining reliable ohmic contact
with the skin. Therefore, a long-term sEMG measurement by the conventional method
causes irritation and discomfort and is a potential cause of skin allergy and inﬂammation.
These are considerably disadvantageous for the use of sEMG in human-machineinterfaces because practical interfaces must be as noninvasive and nonintrusive as
possible to gain broad acceptance from ordinary users.
To overcome the above disadvantages, we have focused on capacitive sensing, whichallows the detection of alternating electrical potential through an inserted thin insulator.
Capacitive sensing has been widely applied to measuring electrocardiograms (ECGs).(13–19)
Recent progresses were surveyed in a methodological review by Chi et al.(20) A capacitive
sensing system developed for ECG, however, cannot be simply applied to sEMG becauseof the following reasons. First, the frequency ranges are different between ECG and
EMG. Second, the electrode should be smaller in EMG sensing than in ECG sensing
because a large electrode would easily suffer from interference from adjacent muscles.
In this study, a new capacitive sensing system is developed for sEMG, which enables themeasurement of sEMG through commonly available cloth, and its feasibility is explored.

2.

Capacitive Sensing

Capacitive sensing is based on the principle of the capacitive (or insulated) elec­
trode, which is an electrode requiring no electrolytic paste or conductive adhesive.
Conductive electrodes are capacitively coupled with the skin through an insulator (Fig.
1). The electrodes can carry an alternating bioelectrical current equivalently through thecapacitance of each coupling. In previous studies, various in­sulators, such as anodized
aluminum,(21,22) silicon dioxide,(23,24) pyre varnish,(25) anodic insulated tantalum oxide,(26)and barium titanate(27) were tested because these materials exhibit high permittivity.
Capacitive sensing is facilitated with an impedance-transforming circuit to matchthe high impedance of the coupling to the low impedance required by the subsequentcircuitry. A ﬁeld-effect transistor (FET) source follower has been used as an impedancetransforming circuit, and is sometimes mounted in an electrode housing to reduce thecommon mode interference caused by power lines.(23,25)

3.

Materials and Methods

3.1

Flexible fabric electrodeThe sEMG signal was picked up using three electrodes, each of which was made ofa rectangular sheet (20×50 mm2) of conductive fabric with conductive acrylic adhesive(Kitagawa Industries, CSTK-100). Commonly available cloth, such as woven cotton,

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Capacitive
Couplings
Conductor

Conductor

Insulator(Cloth)Conductive

Capacitor

Capacitor

Skin (Conductor)
Body (Muscle)
Fig. 1. Schematic model of two capacitive electrodes coupled to the skin with inserted cloth andits equivalent circuit.

was substituted for the rigid insulators. This enabled the insulator and electrodes todeform and achieve a closer ﬁt to the coupled region, and accordingly, no electronic
circuit was mounted in the electrode housing. The cloth insulator would also relieve
the irritation, allergy and discomfort experienced with conventional skin-to-electrodecoupling. The electrodes were attached to an insulating support made of ﬂexible
chloroethylene, and a convex connector (diameter: 5 mm) was inserted between the
support and each electrode (Fig. 2). The two lead electrodes at both ends were connected
to a measuring device, to be described in the next section, with shielded wires. The
center electrode was used as a reference electrode to reduce common mode noise as
much as possible. The skin-cloth­ electrode coupling was held together by a rubber band.

3.2

Ampliﬁer unit

An ampliﬁer unit with ﬁltering and ampliﬁcation circuitry was assembled bymodifying the measuring circuit developed for the capacitive sensing of ECG.(13) In what
follows, we will refer to this unit as a measuring device. The sensing circuitry for sEMG
is more sensitive to power line interference because of the differences in the frequency
ranges and the size of the electrodes between ECG and EMG sensing. Two major
modiﬁcations were made to minimize the interference. First, the guarding technique
was employed in the lead wires (550 mm) between an instrumentation ampliﬁer and theelectrodes to reduce the inﬂuence of leakage current. Second, the driven-ground-plane(DGP) technique(28) was introduced to improve the signal-to-noise ratio (S/N). This
technique is a modiﬁcation of the driven-rightleg (DRL) technique,(29) which has been
widely used in ECG measurement. In the DRL technique, the residual of common modenoise between the two differential inputs is inverted and directly fed back to the body
earth electrode. Several groups have demonstrated its effect in reducing common modenoise.(15,19,30,31) Similarly, in the DGP technique, the feedback is accomplished via anelectrically noncontact reference electrode. The center electrode in Fig. 2 was used forthis feedback.

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Sensors and Materials, Vol. 24, No. 6 (2012)lead electrode

20mm

50mm
110mm

referenceelectrodelead electrode

90mm
Fig. 2. Image of the measuring electrode unit made of conductive fabric and mounted on a
ﬂexible insulating support.

Figure 3 shows the electronic circuitry and element values of the measuring device.
The device consists of two voltage followers, an instrumentation ampliﬁer, a high-passﬁlter, a notch ﬁlter (NF Corporation, SR-2BE), a low-pass ﬁlter and three inverting
ampliﬁers. As shown in Fig. 3, the voltage followers are incorporated at the front-end
in the instrumentation ampliﬁer IC. These voltage followers function as the impedancetransforming circuit to match the large coupling impedance that is due to the small
dielectric constant of the inserted cloth. The actual input impedance of the IC on asubstrate was 1.0 GΩ at 1.0 Hz. The circuit constants in the high- and low-pass ﬁlters
were set to obtain a cutoff frequency of 20 and 1,000 Hz, respectively. The notch ﬁlter
was used to reduce the 50 Hz power line interference. The impedance-transforming
circuit was not mounted in the electrode housing so as to avoid a hard contact with the
skin.

3.3

Measurement of frequency characteristics of the measuring device

The frequency characteristics of the developed measuring device were investigatedusing synthetic skin. The experimental setup is shown in Fig. 4. Measuring electrodes
having identical areas of 1,000 mm2 were placed on an insulating support and covered
with a cotton sheet 350 µm thick. Then, three sheets of synthetic skin made of squarepieces of conductive fabric having the same area as the measuring electrode wereattached to the cloth at locations just above each of the measuring electrodes, so thatthe whole area of each measuring electrode was involved in capacitive coupling witheach sheet of synthetic skin. Using a weight, a pressure of 424 Pa was applied to theelectrodes, the cloth and the synthetic skin. This condition was selected so that it wasslightly lower than the mean pressure exerted by the rubber band. A sinusoidal wave

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Sensors and Materials, Vol. 24, No. 6 (2012)
Instrumentation amplifier IC
ShieldElectrode 1
ShieldOver-Voltage
Protection

+

+1

–

R1

R1
ShieldElectrode 2
Shield–Over-Voltage
Protection

High-pass Filter

R3

C2
R7

R2

–
+

R2

R3

R3

Butterworth
Notch Filter(SR-2BE,NF Corp.)fo: 50Hz

C1

C1

–
+

R6

Inverting amplifier

R5

C2

C2

C4

–

+

R8

R10

R10

–

Inverting amplifier

R11
R4

–

R10
R9

C3

+

+

+1

Driven Reference Electrode

Electrode 3(Reference)

Low-pass Filter

C1

R3

R4

–

–

+

+

R13
R4

+
Capacitance [F]Resistance [Ω]
R1

500k

R9

30.88k

C

R2

25k

R10

216.3k

C

1µ

R3

60k

R11

40k

C3

0.1µ

R4

1k

R12

975k

C

0.033µ

R5

100k

R13

25k

R6

7.28k

R14

961k

R7

5.29k

R15

2k

R8

17.85kR16

833k

Inverting amplifier

–

R15
R4

+

R12
R14

1µ

–
+

R16

Fig. 3. Electronic circuitry and element values of the measuring device.

WeightSynthetic body surfaces
Cotton cloth (350 µm thick)

Lead electrodes
Reference electrodeFig. 4. Experimental setup for measuring the frequency response of the developed device.

was input from an oscillator to the synthetic skin and the output from the device wasmeasured. This combination of inputting and measurement was conducted repeatedly
within a frequency range from 1 to 5,000 Hz. A frequency response without any cloth

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Sensors and Materials, Vol. 24, No. 6 (2012)

inserted was also measured for the same frequency range. In addition, S/N and CMRR
with and without the cloth were measured at 100 Hz using the same setup. The sEMG
obtained without the cloth corresponded to the conventional pasteless, direct-contact
measurement. The room temperature was 22°C and the relative humidity was 38%
during the measurements.

3.4

Comparison to a commercial device

3.5

Comparison of mean power spectra

3.6

Inﬂuence of relative humidity on the signal

To compare sEMG signals obtained by the proposed and conventional methods, asimultaneous measurement was performed. One male volunteer, aged 21, participatedin the measurement. The surface of the right brachium of the subject was covered
with a cotton cloth 350 µm thick. The unit of ﬂexible electrodes was clamped on theright biceps brachii muscle over the cloth. An sEMG signal was recorded using thedeveloped device from the site while the subject repeated isometrical contraction andrelaxation of the muscles. A directly measured sEMG signal was recorded using acommercial bioampliﬁer (Teac Instruments, BA1104M) and a commercial telemeter
unit (Teac Instruments, TU-4). Two disposable electrodes (Nihon Kohden, VitrodeF-150M) were attached directly to the skin at both sides of the ﬂexible electrode unit
along the muscle. A body earth electrode for conventional measurement was placed at
the left elbow. The sEMG signals obtained with the developed and commercial devices
were simultaneously digitized at intervals of 100 µs with 16-bit resolution using an A/D
converter and stored in a personal computer using a data acquisition system (Biopac
Systems, MP-150 system). The bandwidth of the commercial device was 5.3–1,000 Hz.
Another measurement was conducted for the rectus femoris muscle of the left thigh in asimilar way. The room temperature was 23°C and the relative humidity was 24% during
the measurements.
The volunteer for the simultaneous measurement also participated in anotherexperiment to measure the power spectra. First, an sEMG signal from the left rectus
femoris muscle was measured through the 350-µm-thick cloth using the developed
device while the subject repeated contraction and relaxation ten times. Then, an sEMG
signal from the same site, but without any cloth, was measured using the same device
during ten repetitive contractions. After the measurements, 10 segments (800 ms) of thesEMG signal were selected from each recording, and a power spectrum was calculated
for each segment by FFT. Finally, the mean power spectrum was computed from the
10 spectra for each recording. The power spectra with and without the inserted cloth
were compared. The latter corresponds to the conventional pasteless, direct-contact
approach. The room temperature was 25°C and the relative humidity was 24% during
the measurements.
The relative humidity of the room air is considered to inﬂuence the signal quality.
Accordingly, this inﬂuence was investigated. The impedance of the inserted cotton wasmeasured with a precision LCR meter (Agilent Technologies, 428A) at different relative341

Sensors and Materials, Vol. 24, No. 6 (2012)

humidities (21, 30, 40, and 51%). Then, the measured data were ﬁtted by|Zcloth(ƒ)| =(Rcont+Rcloth)2 + (2πƒCclothRcontRcloth)21+(2πƒCclothRcloth)2,(1)

where Ccloth and Rcloth are the capacitance and DC resistance of the cloth, respectively,
Rcont is the contact resistance between the cloth and the conductive surfaces, and f is thefrequency. The equation was derived using the model of the cloth shown in Fig. 5.

4.

Results

4.1

Frequency characteristics of the measuring device

4.2

Comparison to commercial device

The frequency-gain response without any inserted cloth (labeled “Conventional”)
showed a ﬂat gain of about 66 dB within the frequency range between 20 and 1,000
Hz, except at 50 Hz, as designed (Fig. 6). The CMRR was 67 dB (Table 1). These
characteristics meet the requirements of the Japanese Industrial Standard (JIS) for
electromyography, and thus, the developed device can be used to measure an ordinary
EMG when the electrode unit is attached directly to the skin. With the 350-µm-thickcloth inserted (labeled “Proposed” in Fig. 6), the frequency-gain response had almost a
ﬂat gain of around 56 dB within the frequency range between 20 and 1,000 Hz.
Figure 7 shows typical recordings obtained from the biceps brachii muscle andthe rectus femoris muscle. The output signal of the developed device showed ﬁrings
synchronized with the sEMG obtained with the commercial device. The waveforms,
however, were not completely identical because of a slight difference in the detection
site. The greater noise was probably due to the higher input impedance of the developed
system.

Zcloth
CclothRcont

RclothFig. 5. Equivalent circuit of the cloth. Rcont is the contact resistance between the cloth andconductive surfaces such as the electrode or skin. Rcloth and Ccloth are the DC resistance andcapacitance of the cloth, respectively.

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Sensors and Materials, Vol. 24, No. 6 (2012)

70
50
30
10
Conventional-10

Proposed
-30

1

10

100
1000
Frequency [Hz]10000

Fig. 6. Frequency-gain responses measured directly from synthetic skin (labeled “Conventional”)
and with the 350-µm-thick cloth inserted (labeled “Proposed”). Grey and black dashed horizontal
lines indicate −3 dB from the maximum gain under each condition. The gain at 50 Hz was sharply
attenuated because a notch ﬁlter was used in the device.
Table 1
Speciﬁcations of the developed device under conditions with and without the 350-µm-thick cotton
cloth. The room temperature was 22°C and the relative humidity was 35% during the CMRR
measurement.
Gain
Pass band
S/N at 100 Hz
CMRR at 100 Hz

Without cloth (direct) With cotton cloth (424 Pa)
66 dB
56 dB
20–1,000 Hz
20–1,000 Hz
15.4 dB
9.0 dB
67 dB
38 dB

4.3

Comparison of mean power spectra

4.4

Inﬂuence of relative humidity on the signal

Figure 8 shows the mean power spectra obtained with and without cloth insertion
using the developed device. At frequencies between 20 and 500 Hz, no notabledifferences were observed between the two power spectra, except for increases in power
at harmonic frequencies of the 50 Hz power line, such as 250, 350, and 450 Hz. Thus,
alteration of the frequency components in the obtained sEMG resulting from cloth
insertion seemed minimal in this frequency range. In contrast, different slopes wereobserved in a higher frequency band above 500 Hz.
The frequency characteristics of the impedance of the inserted cloth are shown in Fig. 9.
Black circles represent the measured impedance at the humidity of 21%. Similarly, the343

Sensors and Materials, Vol. 24, No. 6 (2012)
3.00

0.00

mV

EMG (Proposed)

1.50

-1.50

EMG (Conventional)

-3.00
2.00

0.00
-1.00
seconds

12.0

mV

1.00

27.0

42.0

-2.00
57.0

(a)

EMG (Proposed)

4.00

0.00

mV

2.00

-2.00

EMG (Conventional)

-4.00
4.00

0.00
-2.00
seconds

43.00

mV

2.00

58.00

73.00

-4.00
88.00

(b)
Fig. 7. Simultaneous recordings from (a) biceps brachii muscle and (b) rectus femoris muscle
using the developed device with a cotton sheet 350 µm thick (top) and using a commercial
electromyograph without any cloth (bottom).

impedance was measured at humidity levels of 30% (grey circles), 40% (black triangles),and 51% (grey triangles). The lines in the graph indicate the best-ﬁt curves obtained byapplying eq. (1) to the measured impedance data. The impedance of the cloth decreases
as the relative humidity increases, particularly in a low-frequency range.

5.

Discussion

5.1

Feasibility of capacitive sensing of sEMG

As shown in Fig. 6, the frequency responses with and without the cloth are similar,
although the S/N and CMRR were reduced by 6.4 and 29 dB, respectively. The
recordings obtained from the biceps brachii and rectus femoris muscles (Fig. 7) showedﬁrings synchronized with those obtained with the commercial device. These results
demonstrate the feasibility of the developed system for detecting sEMG even with the344

Sensors and Materials, Vol. 24, No. 6 (2012)

0.1
0.01
0.001
0.0001

Conventional
Proposed

0.0000110100
Frequency [Hz]1000

Fig. 8. Mean power spectra of sEMG measured from the rectus femoris muscle by the proposedand commercial methods.

250

21%
30%

200

40%

150

51%

100
50
0

1

10
100
Frequency [Hz]1000

Fig. 9. Frequency-impedance responses of the inserted cloth in room air at dif­ferent humidities.
The data were measured with a precision LCR meter (Agilent Technologies, 4284A). Solid lines
in the graph are the best-ﬁt curves obtained by applying eq. (1).

cloth inserted, in particular, for nonclinical, human-machine interface applications. It
should be noted that without the two modiﬁcations from the system for ECG, negligiblesEMG signals were obtained with the cloth inserted (data not shown).

Sensors and Materials, Vol. 24, No. 6 (2012)

345

There are some issues to be addressed in the future. First, the S/N and CMRR remain
to be further improved. Another important issue is the suppression of the harmonics
of the 50 Hz power line in the signal, because the increase in the spectral power inthe frequency range above 500 Hz with the cloth insertion may be caused by theaccumulation of harmonics power.

5.2

Inﬂuence of relative humidity on the signal

5.3

Inﬂuence of perspiration on the signal

6.

Conclusions

As shown in Fig. 9, the relative humidity of the room air inﬂuences the impedanceof the inserted cloth, and thus, the quality of the resulting sEMG signal. The relative
humidity during the simultaneous measurements and the power spectra measurements
was 24%, which was lower than the relative humidity observed under ordinary
circumstances. The S/N of the output signal from the developed system can be improved
if the measurement is conducted in a more ordinary, humid environment.
Geddes and Valentinuzzi(32) reported that dry electrodes were really dry only whenﬁrst applied because of subsequent perspiration, and that contact impedance 20 min after
electrode application became between one-fourth and one-ﬁfth the initial impedance.
During capacitive sensing with a cloth, a similar phenomenon will presumably occur,
particularly in a long-term measurement. Because the decrease in the impedance ofthe cloth reduces the voltage loss at the electrode-cloth-skin coupling, perspiration will
increase the amplitude of the output signal and bring it close to the level measured using
directly attached electrodes. Furthermore, the decline in the impedance of the coupling isthought to reduce the interference caused by power lines. Perspiration will thus improve
the S/N of the output signal. In fact, we have observed an improvement in S/N when the
ECG was capacitively measured during sleep through cloth on the dorsal surface of thesubjects.(13)We have proposed a new approach for obtaining the sEMG under conditions wherebya thin cloth is inserted between the electrodes and the skin. We fabricated a prototype
measuring device incorporating this approach, and validated the proposed approach. The
examination of the developed system yielded the following results.
1. The frequency-gain responses indicated that the measuring system was capable ofdetecting sEMG not only when the electrodes were attached directly to the skin, butalso when a 350-µm-thick cotton cloth was inserted between the electrodes and theskin.
2. The proposed method was able to yield a clearly visible sEMG from the biceps
brachii muscle and rectus femoris muscle even with the cloth inserted, despite the factthat S/N and CMRR deteriorated to 6.4 and 29 dB, respectively, upon cloth insertion.
Because the method eliminates discomfort due to conventional skin-to-electrodecoupling, it appears promising for application to wearable sEMG sensors for humanmachine interfaces, such as Human Adaptive Mechatronics.

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