ChoHyunseung1
YangJinhee1
LeeJeongwhan2
LeeJoohyeon3
-
(Institute of Symbiotic Life-TECH, Yonsei University / 50 Yonsei-ro, Seodaemun-gu,
Seoul, Korea {hyunseung-cho, sjinnie7}@yonsei.ac.kr
)
-
(College of Science and Technology, Konkuk University / 120 Neungdong-ro Gwagjin-gu,
Seoul, Korea jwlee@kku.ac.kr)
-
(Departmet of Clothing & Textiles, Yonsei University / 50 Yonsei-ro, Seodaemun-gu,
Seoul, Korea ljhyeon@yonsei.ac.kr)
Copyright © The Institute of Electronics and Information Engineers(IEIE)
Keywords
Textile electrodes, Configuration method, Heart activity signal sensing performance, Motion artifacts, Wearable platforms
1. Introduction
A major technological trend for the collection and monitoring of individual health
data is heart activity monitoring technology that combines smart textiles, comfortable
clothing design, and biosignal sensing technology. Among the various biodata that
can be measured in the body, electrocardiogram (ECG) signals are biodata that reflect
the heart activity, exhibiting a steady period and regular rhythm under normal conditions.
However, the period and waveform characteristics can change according to changes in
heart activity and the occurrence of lesions. The components of an ECG signal can
be analyzed to detect the presence of heart abnormalities or areas with lesions, and
pathological/ physiological heart abnormalities can be identified to determine heart
disease diagnoses [1]. Therefore, if heart activity signals such as ECG signals can be detected anytime
and anywhere in an unconstrained and unconscious manner, these signals can be extremely
important pieces of an individual’s health data.
Or is "smart" supposed to describe all three? If so, recommend putting "smart"
in front of each of the three for clarity. If not, please specify the intended type
of clothing design.
However, skin-electrode impedance is one of the major problems in biosignal measurements,
because it interferes with accurate signal measurements. When an electrode comes in
contact with the stratum corneum (the outermost skin of the body, composed of dead
cells and oil), it generates significant resistance, and consequently, reduces the
quality of the biosignals [2]. Thus, studies to reduce skin-electrode impedance are being conducted in electrode
development. For example, to measure biosignals, a method was developed in which an
electrode with a spike prepared with micro-processing technology pierces the outermost
skin layer. However, because of the invasive nature of this spike electrode, the user
has to endure both pain and the risk of infection when attaching it [3]. In the present study, because the skin-electrode impedance is reduced as the contact
surface increases, an attempt to minimize the skin-electrode impedance is made by
developing a convex electrode and adjusting the electrode size. The objectives are
to improve the adhesion between skin and electrode, as well as to increase the contact
surface. Based on this, measures to obtain reliable heart activity signals were acquired.
This study focuses on how the textile electrode structure affects heart activity
signal acquisition. Six contact-type textile electrodes manipulated for electrode
size and configuration were fabricated. Heart activity signal-sensing performance
was compared to obtain the textile electrode structural requirements that can minimize
motion artifacts and acquire high-quality signals.
2. Theoretical Background
2.1 Heart Activity Signals
The heart is composed of several types of tissue. The typical cells of each tissue
type exhibit electrical excitability and unique action potential waveforms. In a beating
heart, electrical signals occur, and the electrical potential caused by the heart
appears throughout the torso and body surface. Therefore, potential differences that
occur from heart activity can be measured by attaching electrodes to the body surface
and measuring the voltage between the electrodes. An ECG records the electrical displacement
that occurs during heart activity. In particular, the ECG shows the electrical potential
that measures the electrical activity of the heart’s ventricles, as obtained from
the body’s surface. The major characteristics of a typical ECG signal are P, Q, R,
S, and T waves. P~waves occur due to atrial depolarization. The QRS complex is from
ventricle depolarization. T waves are generated by ventricle repolarization. Atrial
repolarization is masked by the QRS complex and does not appear under normal circumstances.
Among these, the QRS complex (or R wave), which occurs from ventricle depolarization,
is the typical heart activity signal with the largest potential difference. In addition,
a person’s heart-health status and the effects of exercise can be detected through
an ECG and the heart rate signals. This is because the heartbeat is used as an important
body status-monitoring indicator in the field of athletics from a preventative perspective,
and in the field of medicine from a therapeutic perspective [4].
Fig. 1. Different sequences of a typical ECG signal.
2.2 Research and Development Trends in Wearable Heart Activity Signal Sensing Technology
Wearable heart activity signal-sensing platforms generally consist of a textile
electrode that acts as an input interface, a textile signal line, a device (including
a controller, short distance communications, and power supply), and a wearable platform
(band- or clothing-type). Studies on wearable heart activity signal-sensing technology
have been developing rapidly from the early phase in the 2000s until now. During that
early phase, research and development began on a project for the Future Soldier System,
as well as studies on biosignal-monitoring technology by the Philips company (EU).
This era’s wearable heart activity signal-sensing platforms were developed in the
form of various small devices and wires that were simply attached to existing cloth-band
or clothing platforms. An effort that centered on Sensatex, Inc. of the United States
began development of wearable heart activity signal-sensing electrodes and signal
wires that used carbon fibers and metal threads. In the mid-2000s, Germany’s Fraunhofer
research center, the United States’s Textronics, Inc. (which merged with Adidas in
2009), and an EU research project consortium, performed a wider variety of studies
on integrating heart activity sensing electrodes into textiles. Consequently, since
2006, heart activity signal-sensing electrodes and signal wires have advanced to the
level of being developed as fabrics, such as silver- or copper-based metal threads,
carbon fibers, and conductive polymers. Subsequently, since 2010, wearable heart activity
signal-sensing technology has advanced further with the development of heart activity
signal-sensing textile electrodes and signal wires that use materials like grapheme,
silver nanowires, a wider variety of conductive polymers, metal threads based on silver
nanomaterials, etc. Research and development also began on system on textile (SOT)
technology that incorporated the existing controller electronic circuits into fiber
textiles [5-16].
The approach proposed in the previous research was based on applying jacquard
fabric electrodes to a textile electrode when obtaining an ECG. Song et al. (2009)
fabricated textile electrodes by weaving, knitting, or embroidering conductive yarn.
The textile electrode used in biosignal sensing was designed in the form of a jacquard
woven structure to obtain the ECG readings. The proposed textile electrodes were composed
of two groups made up of warps with either unremoved 100% warps or half-removed 50%
warps of jacquard woven electrodes that were convex or flat and that came either with
or without conductive paste. The ECG measurements from the textile electrodes were
taken three times with the patient resting. According to the results, the convex jacquard
electrodes of half-removed 50% warps with conductive paste provided the most significant
SNR improvement (33.67 dB) [8].
Cho et al. (2018) investigated the effect of the contact-type textile electrode
structure on heart activity signal acquisition. In this study, however, we devised
six contact-type textile electrodes where the electrode size and configuration were
manipulated for measuring heart activity signals using computerized embroidery. Heart
activity signals were measured using a modified lead II and by attaching each textile
electrode to a chest band worn by four healthy male subjects in a static standing
posture. The results showed that in the contact-type textile electrodes, the convex
electrode obtained better quality signals than the flat electrode. However, regarding
electrode size, no significant difference was found in the heart signal acquisitions
from three electrode sizes. These results suggest that the configuration method, which
is one of the two requirements of a contact-type textile electrode structure for heart
activity signal acquisition, has a critical effect on the performance of heart activity
signal acquisition for wearable healthcare [15].
Gi et al. in 2015 developed a feasible structure for a textile-based inductive
sensor using a machine embroidery method, and applied it to a noncontact-type vital
sign sensing device based on the principle of magnetically induced conductivity. The
mechanical heart activity signals acquired through the inductive sensor embroidered
with conductive textile on fabric were compared with lead II ECG signals and with
respiration signals, which were simultaneously measured in five subjects. The analysis
showed that the locations of the R-peak in the ECG signal were highly associated with
sharp peaks in the signals obtained through the textile-based inductive sensor (r=0.9681).
The results helped to determine the feasibility of the developed textile-based inductive
sensor as a measurement device for the heart rate and for respiration characteristics
[16].
The original implies the authors of $\textit{this}$ study assessed the results
of the study by Gi et al. If that was intended, however, please reject these changes,
but please clarify how the two different studies are related to each other.
Fig. 2. The weave structure of the jacquard fabric electrodes with 50% of the warp ends removed (Group 2). The electrodes have half of the warp ends removed from the square part of the silver covering yarn [9].
Fig. 3. The embroidered textile-based inductor and its electromagnetic characteristics [15].
2.3 Effects of Textile Electrode Location and Shape on Wearable Heart Activity Signal
Sensing
In the clinical field, silver/silver chloride (Ag/AgCl) disposable metal electrodes
are used to measure ECG signals. However, disposable Ag/AgCl electrodes cannot be
used for long periods because allergic reactions to the electrolyte gel can occur
when the electrodes are attached to the body surface for a long time. Furthermore,
they oxidize easily. Therefore, in wearables that require long-term connection between
the electrodes and the body for heart activity signal monitoring, studies have been
conducted on textile electrodes that can replace Ag/AgCl metal electrodes [5-8, 12,17-26].
Textile electrodes are generally used for measurement of resistance, piezoelectric
voltage, capacitance, and inductive capacity. The most typical form of a wearable
heart activity-sensing textile electrode is the contact-type, which detects heart
activity only when contact is established between the skin and the electrode
[17,18, 22,25-27].
However, signal sensing can be interrupted, or measurement noise can occur, if
a gap appears between the skin and the electrode owing to the wearer’s movement or
changes in the contact point’s location. Heart activity measured using contact-type
textile electrodes is generally based on resistive or piezoelectric measurement principles.
Studies on noncontact-type textile electrodes for heart activity signal-sensing have
included studies on loose-contact heartbeat sensing textile electrodes [24] and radial artery noncontact pulse measurement systems [28] that use the Colpitts oscillator capacitance. The signal-detecting performance from
these methods was found to be relatively low compared to those of contact-type heartbeat
sensing electrodes. The reason radial artery pulse measurement using Colpitts oscillator
capacitance exhibits low performance is that as the electrode’s cross-section area
becomes larger, the electrical field forms within a larger range, thereby generating
more noise that affects the capacitance. This sensing method is limited by the fact
that it reacts sensitively to heart activity signals as well as noise, and measurement
is difficult when the subject is moving. Furthermore, if subjects have thick layers
of fat on their wrists, the detected signals are extremely weak; therefore, it is
difficult to distinguish the signal from noise.
Another principle for heart activity signal-sensing, noncontact-type textile
electrodes is to use the magnetically induced, conductivity-based heartbeat sensing
method. That method is less affected by motion than electrical field-based heartbeat
sensing methods. If a time-varying magnetic field is applied to excite tissues (such
as muscles) in the body, eddy currents occur from the time-varying magnetic field,
and these currents cause reinduced magnetic fields to create changes in the inductance
of the excitation coil. These effects appear as fluctuations in the oscillation frequency,
and are the basis for the magnetically induced conductivity method [29]. As a factor that affects magnetically induced conductivity-based heart activity
sensing, the effect of the textile electrode’s shape characteristics was analyzed
in a study by Gi et al. in 2013 [19]. In addition, a fiber-based inductive sensor structure that uses a computer embroidery
method was developed, and mechanical heart activity signals obtained by this method
were compared to lead II ECG signals. The results indicated that the location of the
R-peaks in the ECG signals exhibited a high correlation (r~=~0.9681) with the peaks
of the signals obtained by the fiber-based inductive sensor.
Fig. 4. Induced textile electrodes in a modified lead II method.
3. Methods and Procedure
3.1 Implementation of Textile Electrodes
The difference in potential caused by electrochemical reactions that occur on
the boundary layer between the body and electrode is called the electrode impedance.
It acts as electrical noise when detecting signals. The contact impedance between
the electrode and the stratum corneum is reduced when the contact surface area is
increased. The aim of this study is to obtain a method for increasing the adhesion
between skin and electrode through the electrode’s structure in order to increase
the contact surface area and reduce electrical noise, such that high-quality signals
can be detected.
In this study, the structures of the textile electrodes are classified according
to two variables: size and configuration. Three electrode sizes were implemented:
1 × 1 (㎝), 2 × 2 (㎝), and 3 × 3 (㎝). Two types of electrode configuration were implemented:
convex electrodes (for which conductive rubber is inserted in the back side of the
electrode and satin embroidery is applied) and flat electrodes (which are embroidered
in a flat manner through a normal satin embroidery method). By combining these three
sizes and two configurations, six heart activity-sensing electrodes were fabricated.
For the electrode material, 280-denier conductive thread (PE silver-plated nickel
mixed thread) was used, with computer embroidery applied to nylon fabric. To measure
heart activity in a modified lead II method, six triangular induced electrode sets
that include one ground (RLD) electrode and two induced electrodes were fabricated
(Fig. 3). The six induced electrode sets were designed to exhibit the same distance to the
electrode center of 2.5 ㎝, even though the electrode structures were different; furthermore,
the wire length was the same.
3.2 Measurement
To verify the performance of the heart activity-measurement electrode, heart
activity-detection experiments were performed on eight males with healthy bodies.
As shown in Fig. 4, the aforementioned contact-type textile electrodes were attached to a chest band,
and heart activity was detected using a modified lead II measurement method. In the
modified lead II method, the center of the RA is positioned on the mammillary line
horizontally, and the center of the ground (RLD) was positioned on the body’s centerline.
To verify the reliability of the heart activity signal acquired by the textile electrodes,
Ag/AgCl electrodes were attached to the body using the standard lead II method to
take measurements and simultaneously provide a reference signal. In the experiment
protocol, measurements were taken for 60 seconds in a standing static posture, and
then for another 60 seconds while walking in place with a heart rate of 80 BPM. To
ensure reliability, the measurements were repeated four times for each subject. A
BIOPAC Systems, Inc. electrocardiogram amplifier was used to collect the ECG signals
from the eight male subjects (Fig. 5).
Fig. 5. Heart activity signal measurement experiment.
Fig. 6. Heart activity signals per second of subject A based on textile electrode type (static state).
3.3 Data Analysis
The BIOPAC ECG100 was used to collect heart activity signals that were detected
in four of the subjects during four rounds of experiments by each textile electrode
type. The signals were sampled at 1~㎑, and the measured raw signal was fed through
a bandpass filter (Butterworth bandpass: order = 10, passing band: 5-15 Hz). To quantitatively
compare the heart activity sensing performance based on the textile electrode structure,
the signal power ratio (SPR) of the heart activity signal detected by each electrode
was calculated. That is, an FFT analysis was conducted, and the signal power ratio
of the 5-15 ㎐ band, in which the heart signal is distributed, was calculated and compared
to the signal power of the entire signal in the frequency space. Using this to normalize
the heart activity signal-detecting results, the interference from noise factors caused
by the difference in the unique characteristics of each subject was controlled. Statistical
analysis that used R was conducted on the calculated SPR value to compare the differences
in heart activity signal-sensing performance based on the six types of textile electrodes:
4. Results and Discussion
The results of the present study were displayed using AcqKnowledge 4.2, and the
detected signals’ waveforms were compared qualitatively. Figs. 6 and 7 show a portion
of the heart activity signal waveform of subject A. The waveform in the top part of
the graph is the reference signal detected by the Ag/AgCl clinical electrode, and
the waveform in the bottom is the heart activity signal detected by the textile electrode
developed in this study. To compare the results that were detected by the six types
of textile electrodes, the size and waveforms of the heart activity detected by each
electrode were analyzed, and the peaks of all the data were found. These results are
shown as a graph in Figs. 6 and 7.
As shown in Figs. 6 and 7, the size of the signal detected by the textile electrodes
tended to be extremely similar to that of the reference signal detected by the Ag/AgCl
electrode. That the textile electrodes developed in this study can detect high-quality
heart activity signals, despite dynamic movements, was verified. In particular, with
the convex electrodes, all three sizes recorded signals with extremely stable waveforms,
and the consistency between the peak points of the detected and reference signals
tended to be extremely high. However, the waveforms of the heart activity signals
detected by the flat electrodes were relatively unstable, compared to those detected
by the convex electrodes.
To quantitatively verify that a difference exists in sensing performance based
on the textile electrode size and configuration method, Matlab was used to analyze
the R-peaks of the heart activity signals detected by the textile and Ag/AgCl electrodes;
furthermore, the SPR value, which is the ratio of the signal component of interest
in the relevant band to the total signal power, was calculated. As shown in Fig. 8, the peak values of the experimental data detected for each subject were found, and
the SPR values were calculated using the data detected from four repeated measurements
for each textile electrode size and configuration method.
The SPR values were used as parameters, and a Kruskal-Wallis test, which is a
nonparametric difference test, was conducted to determine if a difference in performance
was found among the six types of electrode (Table 1). The analysis indicated statistically significant differences among the six electrodes
(p-value = 8.942e-12 [{\textless}0.05]). To confirm that differences existed between
certain combinations of the six electrodes, a Bonferroni post-hoc test was conducted.
In the post-hoc test results, the 1 ${\times}$ 1 (㎝) convex electrodes provided better
performance than the 2 ${\times}$ 2 and 3 ${\times}$ 3 flat electrodes (p-value =
0.00037, p-value = 0.00051). The 2 ${\times}$ 2 convex electrodes provided better
performance than the 1 ${\times}$ 1, 2 ${\times}$ 2, and 3 ${\times}$ 3 flat electrodes
(p-value = 0.01119, p-value = 7.6e-06, p-value = 1.1e-05). Furthermore, the 3 ${\times}$
3 convex electrodes demonstrated better performance than the 1 ${\times}$ 1, 2 ${\times}$
2, and 3 ${\times}$ 3 flat electrodes (p-value = 0.00849, p-value = 5.2e-06, p-value
= 7.6e-06, respectively).
To verify that a difference in performance existed between the textile electrode
configuration methods, a Wilcoxon test, which is a nonparametric t-test (Table 2), was conducted. In the results based on the textile electrode configuration method,
a significant difference was indicated in the SPR values of the heart activity signals
(p-value = 4.831e-14 ({\textless}0.05)). Specifically, the convex electrodes demonstrated
better signal quality than the flat electrodes. This may imply that convex electrodes
are better than flat electrodes at adhering to locations on the skin that are suitable
for heart activity-signal detection, thus increasing the contact surface area, reducing
electrical noise, and detecting higher-quality signals.
To analyze the differences in sensing performance based on electrode size, Kruskal-Wallis
tests were conducted, and the results are shown in Table 3. The performance differences based on the three sizes of textile electrode were analyzed,
and the results indicated no significant difference (p-value = 0.9772 [{\textgreater}0.05]).
Specifically, this study rejects the research hypothesis that when the electrode size
is large, the contact surface area is increased by the same amount, and higher-quality
signals can be detected. This study found that the contact surface area is not increased
simply by increasing the electrode size. When a heart activity signal is detected
by an electrode with a structure that adheres well to a suitable location on the skin,
the contact surface area is increased further, and high-quality signals are detected.
As shown in Table 1, the experiment demonstrated that all three sizes of the convex electrodes (1 ${\times}$
1, 2 ${\times}$ 2, and 3 ${\times}$ 3 ㎝) demonstrated better performance than the
2 ${\times}$ 2 and 3 ${\times}$ 3 flat electrodes; furthermore, these results indicate
that signal quality is affected more by the electrode configuration method than the
electrode size.
In previous research, the sensing performance of textile electrodes was verified
in a static state on four male subjects in their 20s with an average body type. Cho
et al. in 2018 investigated the effect of the contact-type textile electrode structure
on heart activity-signal acquisition. The results showed convex electrodes demonstrated
better signal quality than flat electrodes (p-value = 0.0001367 [<0.05]). However,
the performance differences based on three sizes of textile electrode were analyzed,
and the results indicated no significant difference (p-value = 0.7576 [>0.05]). These
results corresponded to the results of measurement in the dynamic state [15].
From the results above, the configuration method is a textile electrode structure
requirement that significantly affects the quality of contact-type heart activity
signals. Higher-quality heart activity signals can be detected through convex textile
electrodes, regardless of electrode size.
Fig. 7. Heart activity signals per second of subject A based on textile electrode type (dynamic state).
Table 1. Comparison of the six textile electrode types and the post-hoc test.
|
Electrode Type
|
N
|
Mean (SD)
|
Kruskal-Wallis
|
p-value
|
Bonferroni
|
|
1×1 (㎝) convex electrode
|
32
|
-11.62 (2.03)
|
60.644
|
8.947e-12*
|
1×1 (㎝) flat < 2×2 (㎝) convex
1×1 (㎝) flat < 3×3 (㎝) convex
2×2 (㎝) flat < 1×1 (㎝) convex
2×2 (㎝) flat < 2×2 (㎝) convex
2×2 (㎝) flat < 3×3 (㎝) convex
3×3 (㎝) flat < 1×1 (㎝) convex
3×3 (㎝) flat < 2×2 (㎝) convex
3×3 (㎝) flat < 3×3 (㎝) convex
|
|
1×1 (㎝) flat electrode
|
32
|
-14.28 (7.07)
|
|
2×2 (㎝) convex electrode
|
32
|
-10.70 (1.83)
|
|
2×2 (㎝) flat electrode
|
32
|
-16.14 (5.06)
|
|
3×3 (㎝) convex electrode
|
32
|
-10.62 (2.33)
|
|
3×3 (㎝) flat electrode
|
32
|
-16.06 (3.99)
|
* p-value < 0.05
Table 2. Comparison of the textile electrodes based on configuration.
|
Electrode Configuration
|
N
|
Mean (SD)
|
Wilcoxon test
|
p-value
|
|
convex electrode
|
96
|
-10.98 (2.10)
|
7510
|
4.831e-14*
|
|
flat electrode
|
96
|
-15.49 (5.54)
|
* p-value < 0.05
Table 3. Comparison of the textile electrodes based on the three manipulated sizes.
|
Electrode Size
|
N
|
Mean (SD)
|
Kruskal-Wallis
|
p-value
|
Bonferroni
|
|
1×1 (㎝)
|
64
|
-12.95 (5.33)
|
0.0462
|
0.9772
|
-
|
|
2×2 (㎝)
|
64
|
-13.42 (4.67)
|
|
3×3 (㎝)
|
64
|
-13.34 (4.24)
|
* p-value < 0.05
5. Conclusion
Six contact-type textile electrodes for detecting heart activity signals were
fabricated in this study and compared based on the electrode size and configuration
method. Through experiments on human subjects, textile electrode structural requirements
that were more suitable for detecting heart activity signals were discovered. The
results indicate that the electrode configuration method is the structural requirement
that significantly affects the quality of heart activity signals.
In this study, convex contact-type textile electrodes could detect higher-quality
heart activity signals than flat electrodes, and these results matched those of a
study by Song et al. [8]. Furthermore, these results confirm that a convex structure can reduce the electrode
impedance in a contact-type textile electrode. In addition to the electrode configuration,
a size variable was added to analyze the structural requirements of textile electrodes
detecting high-quality signals; however, no difference in sensing performance was
found based on size. However, previous results have consistently shown that the major
factor affecting heart activity signal quality is the textile electrode’s convex configuration,
regardless of the electrode size. This suggests that the textile electrode variables
that must be emphasized in follow-up studies on textile electrode structures are the
requirements related to electrode configuration method rather than electrode size.
The limitation of this study is that the analysis was conducted using data obtained
from a small group of eight subjects, and caution must be taken in generalizing the
results. We believe the results need to be verified in further studies by conducting
experiments on a comparatively larger number of subjects with more diverse body types.
Concurrently, there are plans for implementing a wearable platform with integrated
convex textile electrodes based on this study. Performance improvements in order to
develop smart clothing technology that can detect high-quality heart activity signals
without time or space limitations will be examined in further studies. The heart activity
signal-detecting technology developed through studies that combine smart clothing
with biosignal measurements will be used as an important index for monitoring physical
condition in the preventative and therapeutic medicine fields for customized healthcare.
ACKNOWLEDGMENTS
This research was supported by the Basic Science Research Program through the
National Research Foundation of Korea (NRF) funded by the Ministry of Education (NRF-2020R1A2C1013570
/ NRF-2018R1D1 A1B07049804).
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Author
Hyun-Seung Cho received her Ph.D. degree in Cognitive Science from Yonsei University.
Her research interests are smart clothing, smart healthcare, and wearable sensing
system.
Jin-Hee Yang received her Ph.D. degree in Clothing & Textiles from Yonsei University.
Her research interests are smart clothing, energy harvesting, and wearable sensing
system.
Jeong-Whan Lee He received the B.S. and M.S. degrees in electrical engi- neering
in 1992 and 1994, respectively, and the Ph.D degree in electrical and computer engineering
from the Yonsei University, Seoul, Korea, in 2000. He worked as a member of the technical
staff, M_Application Project Team and a Project Manager, Ubiquitous-health Project
Team at Samsung Advanced Institute of Technology, Kyounggi, Korea, from 2000 to 2004.
Since September 2004, he has been a Faculty Member of the School of Biomedical Engineering,
Konkuk University, Chungju, Korea. His research interests include biomedical signal
processing and instrumentation, wearable/wireless devices for pervasive healthcare,
sensing of vital signs/signatures, heart-rate variability, and biological signal measurements
by electro- magnetic waves and microwave radio-thermometry.
Joo-Hyeon Lee She received the B.S. and M.S. degrees in Clothing & Textile in 1983
and 1985, respectively, and the Ph.D degree in Clothing & Textile from the Yonsei
University, Seoul, Korea, in 1990. She received the A.A.S from Parsons School of Design
in 1992, New York, U.S.A. She worked as a Design Associate of He-Ro group in New York
City, U.S.A. Since 1995, she has been a Faculty Member in Clothing & Textile of the
School of Human ecology at Yonsei University, Seoul, Korea. Her research interests
smart clothing and wearable technology.