DOI: 10.1002/mds3.10016
ORIGINAL ARTICLE
Conductive textile as wearable electrode in intrabody communications Clement O. Asogwa1 | Vlad Libeson2 | Daniel T. H. Lai3 1
College of Engineering & Science, Smart Electronics Systems Research Group, Victoria University, Melbourne, Victoria, Australia
2
Research Director, Ventou Sports and Garments Technology, Melbourne, Australia
3
Smart Electronics Systems Research Group and VU Defence Science Institute (DSI) Liaison, Victoria University, Melbourne, Victoria, Australia Correspondence Clement O. Asogwa, College of Engineering & Science, Smart Electronics Systems Research Group, Victoria University, Melbourne, Vic, Australia. Email:
[email protected]
Abstract Conventional electrodes are widely used for low‐frequency applications such as biopo‐ tential measurements. However, with current advances in intrabody communication (IBC), wearable devices for health monitoring have become popular in biomedical sens‐ ing, data acquisition and high‐frequency communication systems. Wearable dry elec‐ trodes for health monitoring, made from conductive polymers, are often not properly tested for motion artefacts and lack adequate documentation. Traditionally, wet medi‐ cal electrodes made with conductive gel degrades as absorption, and evaporation oc‐ curs. In this work, we introduce a new wearable dry surface electrode from conductive yarn and characterise it by motion artefact effects at 900 kHz in relation to popular wet electrodes. We also compare measurements of skin‐electrode contact impedance and scattering parameters and show that for 60 min of monitoring, dry electrode made of silver conductive textile changed by 0.82% while Ambu WhiteSensor 4500M changed by 2.25% and Noraxon #270 changed by 1.61% at low frequencies. Thus, our wearable dry electrode has more stable skin‐electrode impedance. Our wearable band electrode had a maximum of 1.0 dB loss when a 0 dB signal is transmitted across a 20‐cm channel length in comparison with Ambu WhiteSensor and Noraxon #270 elec‐ trodes which dropped by 2.62 and 2.18 dB, respectively. The result indicates that dry wearable electrodes are more suitable for measurements where body movement is restricted. KEYWORDS
conductive textiles, electrodes, intrabody communication, physiological measurements, sensors
1 | I NTRO D U C TI O N
Usually, medical diagnosis and pathological assessment involves monitoring physiological signs. Other signs monitored by individu‐
Increase in demand for quality health services at reduced cost has
als, sports physiologists and clinicians include changes in an individ‐
led to development in devices that assist remote health monitor‐
ual’s bone mass, hydration state, body fat and so on in relation to
ing without obstructing user’s activities of daily living. In effect,
a recommended baseline. A main advantage of this new system is
real‐time health monitoring technologies are breakthroughs in
the ease and reduced cost of monitoring and follow‐up. Out‐patient
healthcare system (Lai, Begg, & Palaniswami, 2011). Technological
monitoring has caused significant decrease in the high cost of pro‐
developments have brought about dramatic changes in patient’s
longed hospital staying (Park & Jayaraman, 2010), enhanced close
diagnostic methods and monitoring locations from intensive units
monitoring capability and improved early diagnosis and timely alert
in hospitals to individual homes, care centres and respite places.
on patient’s conditions (Custodio, Herrera, López, & Moreno, 2012).
Med Devices Sens. 2018;e10016. https://doi.org/10.1002/mds3.10016
wileyonlinelibrary.com/journal/mds3
© 2018 Wiley Periodicals, Inc. | 1 of 9
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Under various circumstances where sensors are applied the contact
Adaptable versions of dry electrodes that are more flexible to the
interface between the sensors and the human body are electrodes.
curvature of the body like rubber electrode consisting of conductive
Sensitive electrodes and electrodes with separately attached sen‐
polymers and textile has been developed (Gargiulo et al., 2008) with
sors work together to pick up signals from the body, which are trans‐
prospects for wider applications (Ishijima, 1993). Conductive yarn is
mitted to an external device. Interface electrodes are either dry or
the most commonly used material for making flexible dry electrodes.
wet, polarizable and non‐polarizable. Polarizable electrodes from
Conductive textiles are clothing materials that are made to exhibit
metals are characterised by their poor electrode performance due
electrical conductivity. They are used in a wide range of applications
to offset charge distribution around the electrode contact point.
such as electromagnetic shielding, hygienic clothing, sport and fitness
Non‐polarizable electrodes are less susceptible to movement er‐
monitoring, consumer electronics and defence applications (Stoppa
rors and are less affected by offset charge distribution. The most
& Chiolerio, 2014). Conductive fabrics are flexible and conform to
commonly used medical electrode is the silver/silver chloride elec‐
body contours thereby keeping the sensor in the same position and
trode because it is non‐polarizable and exhibits less electrical noise
preventing misplacement. Electrically conductive textiles are made
(Neuman, 2000). Usually gelled electrodes are replaced after some
by spinning, plying or coating fibre thread or yarn with conductive
minutes of monitoring due to skin irritation, allergic reactions and
materials. Recent research shows that conductive fabrics can be used
redness of the skin contact area. Measurement results are also af‐
as electrodes for electrocardiogram (ECG) monitoring (Scilingo et al.,
fected by problems of absorption and evaporation. Consequently,
2005; Stoppa & Chiolerio, 2014), electromyography (EMG) sensing
continuous effort is still made to improve on current performance as
(Custodio et al., 2012) and electroencephalography (EEG) (Lfhede,
well as to adapt dry electrodes as a preferred alternative (Meziane,
Seoane, & Thordstein, 2012). Dry electrodes are mostly used capaci‐
Webster, Attari, & Nimunkar, 2013). Dry electrodes also have flaws.
tively, with partial contact with the body or as a resistive load with
The electrode‐skin impedance of dry electrodes fluctuate after a
complete electrode‐skin contact. When used capacitively, they ex‐
layer of sweat had formed between the electrode contact interface
hibit unstable dielectric property with large base line drifting due to
and the skin depending on the contact pressure, user activities and
motion artefacts (De Luca, Le Fever, & Stulen, 1979) and easily loose
individual sweat development rate. Ideally, electrodes are charac‐
contact with the skin (Valchinov & Pallikarakis, 2004). Several com‐
terised by their biocompatibility, electrical performance, adhesive
mercially available dry electrodes are made with microscale tips of
performance, good signal quality and shelf life (Date & Date, 2015;
silicon wafer with conductive coating (Hsu, Tung, Kuo, & Yang, 2014;
Striefel, 2007). The definition of these recommendations describe
Ng et al., 2009) or nanoscale needles made of nanotubes (Ruffini et al.,
issues relating to electrodes’ electrical properties, end user comfort,
2008). The needle‐like tips are to improve grip on the electrode–skin
minimal interference with daily routine, good contact quality, ease of
contact interface. There are also dry electrodes which enable firm
replacement and the ability not to compromise the integrity of the
contact without skin penetration such as polymer foam electrodes
output signal and the skin when used for either short or long‐term
and conductive textile electrodes (Lin et al., 2011; Shyamkumar et
periods.
al., 2014). Hybrid architecture of capacitive/resistive electrodes with high impedance output suitable for brain computer interface (BCI) and
2 | D RY E LEC TRO D E S
locations with hair are produced (Matthews et al., 2007; Matthews, McDonald, Fridman, Hervirux, & Nielsen da Silva, 2005). Versions suitable for locations with hair are designed for higher electrode im‐
Dry electrodes that offer reduced experimental preparation time do
pedance between 100 and 2,000 kΩ. A more comprehensive study
not require gel or technical expertise to use and can extend length of
on commercial dry electrodes, description and evaluation can be
application with minimal interference to user’s comfort is a challenge
found in the review work of Lopez‐Gordo, Sanchez‐Morillo, and Valle
to researchers (Searle & Kirkup, 2000; Valchinov & Pallikarakis, 2004).
(2014). Skin penetrating electrodes have the advantage of improving
To address these problems and achieve significant success over some
the skin‐electrode impedance but are expensive, prone to infection,
operational aspects of wet electrodes, dry electrodes are designed to
skin irritation and the materials have problems of biological compat‐
operate without artificial electrolyte or gel and are often held in posi‐
ibility with living organisms. Our wearable dry electrode is non‐pen‐
tion with elastic bands rather than adhesives. However, when held
etrating with adjustable elastic band, which makes firm grip on the
against the skin, a layer of sweat usually forms under the electrode
body.
surface in contact with the skin and affects conductivity. In 2004,
The purpose of this work was to introduce a wearable surface
Valchinov and Pallikarakis (2004) predicted that dry electrodes can be
electrode with textile character, using conductive fibre adaptable to
designed to perform just as well as wet electrodes. Three years later,
the curvature of the body with reasonable assurance of safety and
Fonseca et al. (2007) developed a dry electrode that offered consid‐
effectiveness. Our goal is to make dry electrode from conductive
erable advantage over wet electrodes but the reliability test did not
fabric that allows full body contact for clinical studies at frequen‐
include performance under different body movements. (Gruetzmann,
cies associated with intrabody signal communication (IBC) (Asogwa,
Hansen, and Müller (2007) suggested that coating the contact inter‐
Collins, Mclaughlin, & Lai, 2016). The specifications for IBC allow for
face of dry electrodes with conductive foam may improve stability and
data transmission using human tissue as the communication chan‐
resistance to movement artefact in comparison with wet electrodes.
nel. Body movement and exercise causes sweat production, gel
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evaporation and possible absorption through the skin, which influ‐ ences the electrode‐skin impedance and ultimately affects the qual‐ ity of the output signal (Scott, 2003). Our conductive yarn was made by alternate combination of poly‐ mer yarn and metallic yarn to enable uniformity of conduction. The rest of the study was organised as following: section III describes the method we used to make our band electrode and the various tests used to characterise its performance, section IV is the results, section V the discussion and the conclusion in section VI. F I G U R E 1 Schematic for fabricating our band electrode from silver conductive yarn
3 | M E TH O D We used conductive yarns for our band electrode. The conductive yarn is a proprietary product of Ventou sports and garment technology Brooklyn, Melbourne, Australia. Other materials include base fabric, polymer film, elastic band and studs. Table 1 is the electrical properties of the materials. With the conductive fabric, our focus is on achiev‐ ing stability between electrode interface and resistance to movement artefacts (Meziane et al., 2013) and flexibility to the curvature of the body. Other benefits include its biological inertness, cost‐effective‐ ness and availability. We coated the silver conductive fabric with a polymer film to avoid moisture absorption and to separate sweat from bridging the electrodes during use. We then glued them together at 150°C for 45 s that enabled the glue penetrate into the fabric. The polymer‐coated fabric was cut at 1.2 cm diameter and placed on a 1‐ cm perforated space cut out on a non‐conductive fabric as shown in Figure 1. It is heat glued at 150°C for 45 s. The contact electrodes are placed at 4 cm inter‐electrode distance to match the inter‐electrode separation reported in (Asogwa, Collins, et al., 2016) which is based on the closest possible inter‐electrode distance with the single use, Noraxon electrode, model number 270 and Ambu white sensor, Model number 4500M. Figure 1 is a schematic diagram of the production
F I G U R E 2 Wearable band electrodes from silver conductive yarn
process for our locally fabricated band electrode. We fixed a silver me‐ tallic stud on top of the conductive fabric with adjustable band sown to the edge of the insulating material to make a firm wearable elec‐
performance against existing medical electrodes, Noraxon and Ambu
trode as shown in Figure 2.
by measuring the skin‐electrode impedance and compare the changes in contact impedance for 100 min of measurement. We also evaluated
3.1 | Assessment of electrode performance
the effects of motion artefacts using S‐parameters on transmit and re‐
Signal amplitude from bioelectrical measurements is affected by chang‐
same frequency and keeping the same transmission distance.
ceive gains during body movement and without body movement at the
ing skin‐electrode impedance due to moisture absorption and evapo‐
We conducted three experiments for our analysis. First, we
ration, duration of use, frequency, distance between the electrodes
tested the skin‐electrode impedance with the three electrode types
and movement artefacts (Scott, 2003). We characterise our electrode
and measured the S‐parameters of the electrodes when there is body movement and when the body was still. Finally, we tested the
TA B L E 1 Electrical properties of the material
performance of the electrodes after it has been used, washed and
Material
conductive element
Resistivity (Ω/m)
Conductive fabric
Silver
2.0
Base fabric
Non‐conductor
—
Polymer film
Non‐conductor
—
Elastic band
Elastic fabric
—
Stud
Silver base
0.2
dried naturally. 1. Characterisation by skin contact impedance: we used a 50 mi‐ croamp Noraxon MultiTester to measure the skin‐electrode impedances at 20, 100 and 200 Hz which is the frequency often used to test skin contact impedance for surface EMG measurements (Neuman, 2000). The test current is within the
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safety guideline recommended by (ICNIRP). To ensure balance
Solutions, baluns (Coaxial RF transformers, FTB‐1‐1+, turns ratio of
in measurements and comparison, we maintained the same
one, manufactured by Mini‐Circuits and frequency range 0.2–
site position on each subject (Ahlbom et al., 1998). The three
500 MHz). Pre‐gelled self‐adhesive Noraxon electrode model
electrodes were placed as shown in Figure 3 to enable us
number 270, manufactured by Noraxon Inc, Ambu WhiteSensor
observe the changes in the skin contact impedance with the
electrode, model number 4500M, by Ambu Inc. and wearable dry
three electrodes within the same period. All participants signed
surface electrode made of silver conductive fabric were used. Each
the consent to participate.
electrode maintained a 20‐cm separation distance between the
2. Evaluation based on s‐parameters: we used a hand held vector net‐
transmit and receive electrodes. The contact interface of all the
work analyser (VNA) to independently measure the s‐parameters
electrodes is 1 cm in diameter. To verify and compare the s‐param‐
of a propagating signal with our Band electrode, Ambu WhiteSensor
eters of our wearable electrode with conventional wet electrodes,
electrode (model number 4500M) and Noraxon electrode (model
we examined data obtained during artefact‐producing movements,
number 270). The VNA was chosen because it has wider dynamic
and those collected when the body is still and artefact motion is
range, which enables it to sweep across many frequencies. The
expected to be as low as possible. Because we are considering the
setup is as shown in Figure 4. The mini Pro VNA has a frequency
electrode as a wearable device, we chose the arm for our testing
range of 100 kHz to 200 MHz, manufactured by Mini Radio
which is commonly used for wearing watches and wristbands. The electrodes are placed in a central position over the muscle along the frontal view of the Ulna (between the olecranon and processus styloideus ulnae) in order to contain possible muscle migration away from electrode position during experimental periods, espe‐ cially when movement is allowed. The electrodes are placed at 20‐ cm transmit distance as proposed by Asogwa, Seyedi, and Lai (2014) and measured from the wrist to the ankle. Five subjects consisting of four males and a female with body mass index (BMI) ranging from 23.0 to 28.0 kg/m2 volunteered for this study. Subjects were allowed 5 min of settling to allow muscle relaxation before the start of the experiment. For each participant, the 20‐cm transmit distance was measured with tape and the exact positions marked, this made it easy to repeat the experiment on the exact spot with different electrodes. We used baluns to electrically iso‐ late the two ports of the VNA from making a loop through the common ground of the two ports. All the subjects were measured
F I G U R E 3 Measurement of skin‐electrode impedance with Band, Noraxon and Ambu electrodes
first without permitting body movements and then with move‐ ments of the arms and body while sitting at the same position. The measurements were carried out at 10:00 a.m. on each occasion. Average room temperature of 25 ± 0.1°C was maintained through‐ out to minimize the effect of temperature on body metabolism and evaporation. The three electrodes were tested on the same spot on each subject with the average room temperature as stated above. It was difficult ensuring the same magnitude of body move‐ ment between participants but individual movements were con‐ trolled by requiring all participants to bounce a hand ball on the ground once every second in a sitting position while the measure‐ ments were taken. Participants were required to practice the exer‐ cise before the start of the experiment. The process was repeated and the transmit and reflection coefficients measured for when the body was as still as humanly possible, and when body move‐ ment was allowed by bouncing a ball on the ground which allowed the hand to swing back and forth in the air, we note that the force applied on the ball may vary. From Figure 4, the incident and reflected waves were mea‐
F I G U R E 4 Measurement of transmission coefficient with one of the electrodes (Ambu WhiteSensor)
sured with a series of directional couplers as shown diagrammati‐ cally in Figure 5. The scattering matrix relates the outgoing signals
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ASOGWA et al.
b1, b2 to the incoming signals a1, a2 that are incident on the two
electrodes, and our dry wearable band electrode using Noraxon
port.
MultiTester. The result is as depicted in Figure 7. The contact im‐ ⎡ b ⎢ 1 ⎢ b ⎣ 2
⎤ ⎡ S ⎥ = ⎢ 11 ⎥ ⎢ S ⎦ ⎣ 21
S12 ⎤ ⎡ a1 ⎥⎢ S22 ⎥⎦ ⎢⎣ a2
pedance decreased over time in all the electrodes. The changes
⎤ ⎥ ⎥ ⎦
in the electrodes after more than an hour of monitoring in the EMG frequency are shown in Table 2. Noraxon and Ambu had the most significant change in skin‐electrode impedance ranging from
S11 is the reflected signal power measured in dB while S21 refers to the transmission coefficient also measured in dB. b S11 = 1 |(a2 = 0) a1
S21 = −20 log10
S11 = −20 log10
(
Vb2 Va1 Vb1 Va1
a 0.82% change in skin‐electrode impedance at 1 MHz with the band electrode while Ambu and Noraxon changed by 1.58% and 1.52%, as shown in table 3. Again, we compared the three
(2)
electrodes based on their scattering parameters using the set up shown in Figure 4. Our assessment will be based on the changes on the transmission coefficient, S21 using Equation (3) with and
We obtain the transmission and refection coefficients in dB (
a near uniform skin‐electrode impedance with maximum varia‐ tion not exceeding 2.20% at 200 Hz. Similar observation shows
(1)
2 S21 = |(a2 = 0) a1
using the relation
2.46% to 35.90% while our wearable band electrode maintained
) )
without motion artefacts. The results are as depicted in Figure 8 based on the transmission coefficient while Figure 9 is the result
(3)
from the reflection coefficient, S11. Tables 4 and 5 show the effects of movement artefact on transmit and receive signals. Dry wearable electrode has up to 9 dB change in
(4)
3. Electrode performance after washing: to determine the performance of the electrodes after washing, we dipped a pair of the band electrode in a litre of water mixed with 2 ml of washing liquid and washed for 3–5 min and dried naturally. Five days were allowed from the day it was washed before reuse. The dry electrodes were again fixed on the subject’s arm, taking care to ensure that the initial position placement of the electrodes was maintained as previ‐ ously recorded. We took repeated measurements of the transmission losses at five‐minute interval to evaluate the degradation of the electrode performance due to washing in contrast to its performance before washing without allowing body movements. Figure 6 shows the appearance of the band electrodes after three washes. All the experiments followed the ethical regulations of Victoria University Human research ethics committee, approval ID HRE 14‐122.
4 | R E S U LT S
F I G U R E 6 Band electrode after washing
Figure 3 is the set up used for measuring the skin‐electrode impedance of Noraxon and Ambu WhiteSensor as medical
F I G U R E 5 Schematic diagram for measuring scattering parameters of a device under test (DUT)
F I G U R E 7 Comparison of electrode‐skin impedance with Ambu, Noraxon and Band electrodes
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TA B L E 2 Change in electrode‐skin impedance after an hour of monitoring at 20, 100 and 200 Hz
on subjects B and A. This is because dry electrodes easily form a layer of sweat due to the contact pressure between the skin and the electrode which causes sweat quicker when the body is under
Change in impedance
active motion. This caused a drift in the signal output which var‐
Subject
Electrode type
20 Hz (%)
100 Hz (%)
200 Hz (%)
A
Ambu
9.47
29.90
2.46
Noraxon
4.20
11.90
0.16
more amount of electrolytic fluid in contact with the body. This can
Band
0.12
0.20
1.05
be absorbed through the skin as the subject sits still for prolonged
Ambu
6.86
24.10
1.41
measurements or evaporate during body movements or exercises.
Noraxon
3.90
10.60
1.16
When it evaporates, evaporative cooling caused the surface under
Band
1.36
0.90
0.51
the skin to cool thereby reducing sweat which potentially lowered
Ambu
8.90
30.90
2.72
the impact of movement artefacts.
Noraxon
5.27
16.10
3.83
Band
0.38
0.50
1.08
Ambu
14.85
35.90
7.33
Noraxon
13.85
35.90
8.24
Band
0.52
1.36
2.20
Ambu
9.95
22.44
2.33
Noraxon
9.19
17.48
4.28
Band
1.26
1.05
1.46
B
C
D
E
ies as the body movement and exercise varies (Neuman, 2000). The Ambu electrode had 2.0 dB change in signal gain resulting from both movement and non‐body movements. The Ambu WhiteSensor has
Figure 6 depicts the appearance of the electrodes after wash‐ ing. Four different measurements were taken to investigate this. The first two measurements were taken before washing, and the last two after the electrodes had been used and washed. During each measurement, the body was made to be as still as possible. Five days were allowed before every other measurement with each measurement lasting for 30 min to verify the repeatability of the data. The procedure was repeated after the electrodes had been used and washed. There is a large variation in the measure‐ ments from different subjects in Figures 8 and 9. This is largely be‐
amplitude due to movement artefacts similar to AmbuWhite Sensor
cause all the subjects have different body mass indices, metabolic
while Noraxon 270 had the least amount of change due to move‐
rates as well as hydration state. Figure 10 depicts the result of the
ment artefact. Similarly, the reflection of the signal from the trans‐
experiment. Again, measurements Afterwash1 and Afterwash2
mit port was more on the Band electrode followed by Ambu before
separated by 5 days ensured that the electrodes were completely
Noraxon. This means that the effect of movement artefacts is more
dry and without moisture. Firstly, the difference observed in
with the Band and Ambu electrodes than with Noraxon.
the baseline data between each day of measurement was due
Table 6 shows the standard deviation of the signal after 30 min
to different hydration states of the subject (Asogwa, Teshome,
of measurement with the three electrodes. The standard deviation
Lai, & Collins, 2016). The signal gain at intervals of 5 min before
of the signal without body movement was less than 1 dB on 80% of
Beforewash1 and Beforewash2 followed similar pattern and was
the subjects measured with our dry band electrode. On the other
consistent which implies that the recorded data are repeatable.
hand, more than 2.23 dB was the standard deviation observed on
After wash, the first measurement (Afterwash1) had changes
80% of subjects measured with Ambu WhiteSensor electrode and
observed only in the baseline data which has been attributed
60 % recorded above 1.29 dB deviation with Noraxon electrode.
to subject’s hydration sate. The results matched those recorded
This implies our new dry band electrode has less variation in signal
before the electrodes were washed. This indicates that the elec‐
gain when the human body movement is made as still as possible
trode would not be critically affected by washing in the first day
than Ambu and Noraxon. Again, the standard deviation of the signal
of using and washing. However, Afterwash2 (the electrodes were
changed completely when the subjects were allowed to move. The
washed twice and dried before reuse), the result shows a notice‐
minimum deviation measured with the dry band electrode is 1.26
able gradual decline in performance. The changes observed in the
on subject D and as high as 5.54 dB on subject E and above 4.20 dB
magnitude of the gain after each interval of 5 min were ±1 dB on
Change in impedance Frequency (kHz)
Ambu (%)
STD
Noraxon (%)
STD
100
1.12
300
2.71
500
2.15
700
1.97
Band (%)
STD
0.99
0.87
1.86
2.56
0.98
1.21
0.99
1.00
1.36
1.41
0.56 1.31
1.47
0.43
0.99
0.26
1.43
1.28
0.69
0.84
900
2.25
0.33
1.61
0.82
0.82
0.97
1,100
1.58
0.31
1.52
0.71
0.82
0.66
TA B L E 3 Average change in electrode‐ skin impedance after an hour of monitoring between 100–1,100 kHz on a subject
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ASOGWA et al.
F I G U R E 8 Transmit loss measurements on Ambu, Noraxon and Band electrodes at 900 kHz. “A” without movement and “B” with induced movement
F I G U R E 9 Reflection loss measurements, “A” without movement and “B” with induced movement on Ambu, Noraxon and Band electrodes at 900 kHz TA B L E 4 Effect of movement artefact on transmit (change in gain)
TA B L E 5 Effect of movement artefact on receive (change in gain)
Subject
Band (dB)
AmbuWhite sensor (dB)
Noraxon 270 (dB)
Subject
Band (dB)
AmbuWhite sensor (dB)
Noraxon 270 (dB)
A
9.0
9.0
4.0
A
2.80
1.78
0.40
B
4.0
6.0
2.0
B
0.22
3.29
0.09
C
7.0
2.0
5.0
C
0.21
0.07
0.33
D
3.5
8.0
6.0
D
0.20
0.42
0.03
E
5.0
0.5
3.0
E
0.56
0.30
0.12
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TA B L E 6 Standard deviation of measurements with and without movement artefacts No movement
With movement
Subject
Band (dB)
AmbuWhite sensor (dB)
Noraxon 270 (dB)
Band (dB)
AmbuWhite sensor (dB)
Noraxon 270 (dB)
A
0.82
2.56
1.61
4.22
1.85
2.53
B
0.77
2.23
2.18
4.86
2.96
1.47
C
0.93
1.23
1.29
2.30
2.13
2.14
D
0.99
2.62
0.72
1.26
2.07
2.91
E
1.11
2.25
0.93
5.54
3.92
1.32
changes more rapidly while measurements with band (textile elec‐ trode) are more predictable with nearly regular consistent deviation from the mean. Although the amount of change decreases as the fre‐ quency increases, in a tetra polar galvanic coupling, two multiplies the impedance between each pair, which has significant effect to the flow of current on the body. A high contact impedance increases at‐ tenuation and lowers the penetration depth of a transmitted signal. The variation on the signal gain occurred on all the subjects when there is body movement. Noraxon electrode single use model number 270 has smaller amount of fluid than Ambu whiteSensor F I G U R E 1 0 Transmit loss measurement on Band electrode before and after washing
4500M. In our previous work (Asogwa et al., 2014; Asogwa, Collins, et al., 2016), we showed that differences in body mass indices and hydration states affect conductivity and signal propagation across the body. Our test experiment shows that for whatever application,
the average, but consistent. This shows that protracted washing
both endogenic and exogenic processes around the body affects re‐
can impact on the performance of the electrodes with increas‐
sults and analysis and predictions are difficult to accurately quantify
ing probability of degradation as washing increases. Conversely,
but best based on trends. Textile electrodes, however, have lesser
the non‐conducting polymer had obvious shrinkage right from the
degree of irregularity when applied for intrabody communication,
first wash and increased as the number of washes increased but
but like regular medical electrodes have to be applied firmly to avoid
this can be countered by changing the base material of the non‐
noise. Our band electrode has lower impedance drift, making it more
conducting polymer.
suitable for biomedical measurements for situations where body movement is not required. In addition, comparison of the electrode performance before and after wash shows the electrode is compar‐
5 | D I S CU S S I O N
atively reusable. Its effective performance decreases as the number of washing increases and unfit after three reuses.
Electrodes provide the necessary interface between sensing devices and the human body. Assessment of human body composition by galvanic coupling relies on the movement of electrical signal across
6 | CO N C LU S I O N
human body tissues. The nature of the electrodes contributes signif‐ icantly to the measurement outcome. We evaluate the performance
In this work, we show a new dry wearable electrode suitable for low‐
of our new band electrode based on two important parameters,
frequency body composition measurement when the body is as still
changing skin‐electrode impedance and scattering parameters.
as humanly possible. This will provide improved result over wet elec‐
Consistency in skin‐electrode impedance is critical for ensuring that
trodes for assessments that does not require physically exercising
measurements reflect intended outcomes especially when evaluat‐
the body and where baseline measurements of physiological states
ing changing body fluid levels or other forms of body physiological
of the body are expected to be under still condition such as pre‐surgi‐
assessment. The dry wearable electrode has lower drift and more
cal body impedance measurements, baseline measurement of tissue
stable skin‐electrode impedance, changing from 0.51% to 1.08%
before an exercise and assessment of wound healing and recovery
across all the subjects. Both Noraxon and Ambu experienced up
to predetermined readings. Under still condition, popular medical
to 35.90% change in skin‐electrode impedance at frequencies be‐
electrodes increase skin contact impedance by as much as 2.25%
tween 20–200 Hz and 1.52%–22.5% at frequencies in the IBC range.
at 900 kHz and the drift in signal gain is also more than 2.60 dB on
From Table 6, we observed that at low frequencies, skin impedance
average. This value is very high especially in applications requiring
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measurements of tissue recovery and wound healing. Our dry wear‐ able electrode has lower drift and more stable skin‐electrode imped‐ ance, regardless of the differences in gender or body mass index. Dry wearable electrodes are good for baseline readings where there is not much body movement. Results also show the electrodes are washable and reusable which makes it more wearable. The elec‐ trodes work better for frequencies around IBC range. Wet EMG and ECG electrodes give better result at 20 and 200 Hz than at 100 Hz. Our future work will examine closely the effects of movement on the skin‐electrode interface for both dry and wet electrodes, and if there is any correlation with body mass index and gender for applica‐ tions, where body motion is not particularly important.
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How to cite this article: Asogwa CO, Libeson V, Lai DTH. Conductive textile as wearable electrode in intrabody communications. Med Devices Sens. 2018;e10016. https://doi. org/10.1002/mds3.10016
