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Sensors and Actuators A 295 (2019) 678–686

Contents lists available at ScienceDirect

Sensors and Actuators A: Physical

journa l homepage: www.e lsev ier .com/ locate /sna

ighly conformable stretchable dry electrodes based on inexpensiveex substrate for long-term biopotential (EMG/ECG) monitoring

eyman Fayyaz Shahandashti a,1, Hamed Pourkheyrollah a,1, Amir Jahanshahi a,∗,assan Ghafoorifard b

Micro Bio Technology Laboratory (MBTechLab), Department of Electrical Engineering, Amirkabir University of Technology (Tehran Polytechnic), Tehran,ranDepartment of Electrical Engineering, Amirkabir University of Technology (Tehran Polytechnic), Tehran, Iran

r t i c l e i n f o

rticle history:eceived 17 April 2019eceived in revised form 4 June 2019ccepted 19 June 2019vailable online 27 June 2019

eywords:earable electronics

ry electrodeiopotentialstretchable electronics

a b s t r a c t

The comfortable, long-term, accurate, and continuous biopotential monitoring is of paramount impor-tance for biomedical applications. In this respect, dry electrodes are the way forward for the nextgeneration of wearable health monitoring devices. However, it is still challenging to record the bio-electrical signals, i.e., electroencephalogram (EEG), electromyogram (EMG), electrocardiogram (ECG), ofpatients during their everyday life. It is mainly because, soft skin-like, at the same time, electrically con-ducting and biocompatible materials have to be integrated with electrical circuits. Dry electrodes havethe potential to be integrated in this manner into a variety of wearable form factors. However, most dryelectrodes are fabricated on rigid substrates, which do not possess sufficient flexibility and conforma-bility to human skin. Novel flexible/stretchable substrates are a topic of research in the state-of-the-artliterature for dry electrodes in pursuit of the wearable biopotential electrodes.

In this work, a novel flexible and wearable dry electrode technology based on polydimethylsiloxane(PDMS) is presented. It has been tested for EMG and ECG signal acquisition as the proof of principle.The performance of the fabricated dry electrodes is investigated by measuring the contact impedancebetween the skin and the electrode. The impedance is compared with the commercially available wetAg/AgCl electrodes. Overall the contact impedance of the dry electrodes is comparable compared tothe wet Ag/AgCl in the frequency range of interest for biopotential electrodes. Experimentally received

signal quality of EMG/ECG with the dry electrode is comparable to the wet electrode as well. The dryelectrodes do not require any skin preparation nor conductive gels. Thanks to the extreme flexibilityof our electrode, it provides more robust contact as well as less skin irritation. The provided resultssuggest that the presented electrodes/technology have the potential to be used in clinical health-careapplications. To support this claim, long term EMG/ECG results are also provided in the manuscript.

© 2019 Elsevier B.V. All rights reserved.

. Introduction

The biosignals, such as EEG from brain, ECG from heart, andMG from muscles, have been widely used for medical diagno-is [1,2]. The ability to accurately monitor biopotential is vital foresearch applications as well [3–5]. Human activities such as brain

ctivity, heart beating, and muscle movement can be directly mea-ured from the low-level ion current, i.e., biosignals, that is presentn different parts of the body. The electrodes are the key compo-

∗ Corresponding author.E-mail address: Amir.Jahanshahi@aut.ac.ir (A. Jahanshahi).

1 Authors contributed equally to this work.

ttps://doi.org/10.1016/j.sna.2019.06.041924-4247/© 2019 Elsevier B.V. All rights reserved.

nent that responsible for the quality of the received signals. Theycan be classified into two types: dry and wet [6]. The electrolyticconductive gel is used in the former in order to reduce the con-tact impedance between the biopotential electrode and epidermislayer of the skin. Ag/AgCl is currently the gold standard in wet elec-trodes due to its accurate signal recording and ease of use. Theyare vastly used in clinical environments for short term biopotentialmeasurements. However, they are generally not suitable for long-term applications. The quality of the signal decays as the electrolytegel dehydrates over time. Additionally, to improve the quality of thesignal, skin preparation is necessary, which could be painful and

time-consuming. Furthermore, the gel can provoke skin irritationand allergic reactions [7,8]. The mentioned limitations hamper theadoption of wet electrodes in wearable biopotential recording sys-

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ems [9,10]. Reusable, flexible, and preferably washable electrodesre preferred in emerging dry electrode applications.

The dry electrodes have been introduced in the literature asn alternative to wet electrodes [11,12]. Long-term recording viairect skin contact - without gel - is realized using these electrodes.hey typically produce little skin reaction as well. Although theselectrodes have traditionally been infamous to provide poor signaluality compared to wet electrodes, comparable signal qualitiesave also been demonstrated in the literature [9,13]. Nevertheless,igid dry electrodes are still relatively uncomfortable to wear, haveigher skin-electrode impedance, and produce large motion arti-

acts due to slipping and presence of hair. Flexible dry electrodeechnology could mitigate the aforementioned problems due toetter skin contact.

Novel materials and technologies have been developed for theabrication of dry electrodes. The examples include capacitive elec-rodes [14], invasive microneedle electrodes [15,16], electrodes inextile [17,18], and conductive polymer electrodes [19,20]. Theorking principle of a dry, non-contact electrode is the capac-

tive coupling between the human skin and the surface of thelectrode. Among the various technologies, conductive textile elec-rodes illustrate excellent conformability. However, they sufferrom relatively low signal-to-noise ratios due to generally poorkin contact. Conductive polymer electrodes are made from mix-ng elastomers, e.g., PDMS, with conductive materials such as silveranowires or carbon nanotubes (CNTs). These electrodes featurexcellent flexibility and stretchability. However, due to their rel-tively low electrical conductivity, motion artifact compromisesheir performance.

Another technology for dry electrodes is having thin-film metalmbedded in the elastomer. These types of electrodes featurexcellent electrical conductivity due to the presence of the metal.urthermore, they illustrate conformability and stretchability as aesult of having elastomer. This technology can be viewed as anxtension to the emerging stretchable electronics technology. It haseceived extensive attention in the literature due to its favorable

echanical and electrical characteristics. Thin skin-like electronics21], wearable consumer electronics [22], soft robotics [23], tactileensors [24] are just a few examples that have been realized in thetate-of-the-art literature. The dry electrodes fabricated using thisechnology can conform to any curvilinear surface of the humanissue [25–29].

To contribute to the state-of-the-art literature in dry elec-rodes, we have proposed stretchable electrodes based on metalracks embedded in PDMS elastomer. Compared to the relativelyomplex fabrication steps in the literature, a relatively straight-orward, at the same time, inexpensive fabrication technology haseen developed. The starting substrate is the inexpensive com-ercially available flex substrate. Stretchable metal electrodes are

upported with polyimide (PI) layer in order to increase the fatigueeliability [32]. For the first time, sheet-type, compared to the spin-n PI layer, has been used in order to decrease the costs whileeeping the fatigue reliability sufficiently high [30–32]. The con-uctivity of the presented dry electrode technology is relativelyigh compared to the thin-film sputtered metals in the litera-ure, due to the relatively high thickness of metal layer of the flexubstrate. Mechanical and electrical reliability of this technologyas been thoroughly addressed in our previous works [32–34].he presented dry electrode technology has been demonstrated

n monitoring surface EMG/ECG signals. The performance of thelectrodes is proved to be on par with the commercially availableet Ag/AgCl electrodes. The immunity to motion artifacts of the

roposed electrodes is higher than wet Ag/AgCl electrodes. Thelectrodes have been successfully tested in long term scenarios asell.

Fig. 1. Schematic diagram of the (a) EMG and (b) ECG dry electrodes are illustrated.The black regions are skin pads which are connected to the external electrical padsthrough spring shaped electrical interconnects.

2. Design and fabrication

2.1. Design of stretchable electrodes

The existing commercial types of wearable medical electron-ics utilize structures and materials that are mechanically rigid.Despite significant developments in electrical performance of thesedevices, the mechanical properties are lagging behind. To achievea more convenient experience for the user, the devices in con-tact with the human body, need to be more conformable. Flexiblematerials provide deformability and conformability on surfaceswith various topologies. Stretchability, however, is a step forward.Stretchable dry electrodes should comply with a large level of strainwithout significant change in electrical performance. Such elec-tronics have the ability to be bent, folded, twisted, compressed,stretched, and even deformed into arbitrary shapes. The progressof stretchable and wearable sensors necessitates innovations inmaterials, designs, and fabrication processes. Stretchable backboneor substrate is the initial requirement. Typically, flexible poly-mers, rubbers, and metal foils are widely used as substrates dueto their relatively high thermal stability, excellent chemical resis-tance, and extreme mechanical flexibility. Polymers represent themost promising platforms for stretchable technologies due to theirinherent low mechanical stiffness. PDMS has been widely usedas a substrate material in flexible and stretchable devices due toits remarkable biocompatibility, stretchability, and chemical inert-ness.

In this work, we have chosen PDMS as the primary material forthe dry electrode. For the conductive part of the electrodes, met-als with low resistivity are necessary. These materials are rigid andcannot stretch due to their relatively high stiffness. However, thestiffness of any layer is decreased by reducing the thickness. Anyrigid material converts to flexible if the thickness is low enough.In this work, the thin-film metal layer is used as the conductingpart. This type of layer is flexible but not stretchable. To achievetotal stretchability, the metal layer is shaped in meander formto behave as a spring. The thin-film metal layer is embedded inPDMS. Supporting a metal layer with a flexible-only polymer canenhance mechanical properties such as cycling fatigue [32]. In thiswork, metal layers are supported by a PI layer. PI exhibits excellentflexibility, very low creep, high tensile strength, and is compat-ible with the microsystems manufacturing processes. The PI asan intermediate layer between metal and PDMS can increase themechanical strength of the electrodes. Inexpensive commerciallyavailable double-sided flex substrates have been used in this workas the starting substrate [32]. In these substrates, the PI layer issandwiched between two Cu layers. Cu layer of the double-sidedflex is used as the main conducting material of the dry electrode. Toincrease the biocompatibility of the dry electrodes, a thin layer ofAu has been placed on top of the Cu to be in contact with the skin.

As a proof of principle, EMG and ECG dry electrodes are fab-ricated using this technology. Fig. 1a and Fig. 1b illustrate theschematic diagram of the EMG and ECG electrodes, respectively.

680 P. Fayyaz Shahandashti et al. / Sensors and Actuators A 295 (2019) 678–686

F ically. The electrodes are placed on the bicep and arms for EMG and ECG measurements,r mplification, filtering and analog to digital conversion.

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Fig. 3. Complete fabrication steps. a) The starting substrate is a double-sided flex;25 �m thick PI sheet sandwiched between 35 �m thick Cu layers (Pyralux® AP9111R). b) 6 �m thick Au is electroplated on both sides of the Cu layer. c) Strippingaway Au and Cu from the backside of the sample. d–e) Standard photolithographyof Au and Cu layer to pattern electrodes, interconnects and electrical pads. The glasssubstrate is used for handling purposes. f) peeling off the sample from the glass

ig. 2. Biopotential recording setup using the dry electrodes is illustrated schematespectively. The signal that is received with the dry electrodes is sent to PC after a

he EMG electrode has 2 pads with 100 mm2 surface area. Four padsre designed for the ECG electrode with approximately 30 mm2

urface area (Fig. 1). The electrical interconnects are designed ineander form to increase the stretchability of the device. The

esign of the EMG electrodes is adopted using the recommenda-ions of the SENIAM for sEMG sensors [35].

Fig. 2 illustrates the biopotential recording setup schematicallyn this work. The raw signal is pre-amplified, and its bandwidth isimited using a bandpass filter to limit the input noise. It is thenmplified again and is fed to an analog to digital converter. Theigital signal is sent to the PC using a USB port. All the neededlectrical circuitry is designed and made in-house.

.2. Fabrication of stretchable electrodes

The fabrication process of the stretchable dry electrodes isllustrated schematically in Fig. 3. The starting substrate is aouble-sided flex substrate (Pyralux® AP 9111) which consists of au/PI/Cu, where the thickness of Cu and PI are 35 �m and 25 �m,espectively (Fig. 3a). Spin-on PI thickness is defined by the speednd duration of spinning. It also has a smoother surface comparedo the PI in double-sided flex. However, in comparison, the lat-er is remarkably inexpensive and eliminates several processes ofpinning and etching.

The first step is electroplating 6 �m of Au layer on both sidesf the substrate (Fig. 3b). The electroplating is done using a com-ercial, industrial Au plating Cyanide bath [36]. Au and Cu layers

re then stripped away from one side of the sample using Edinburg40% FeCl3) [33] and Aqua Regia (nitric acid and hydrochloric acidn 1:3 mixing ratio) [37] etchant (Fig. 3c). The detailed procedure isimilar to Fig. 3e, as explained in the following. During this process,he other side of the sample is protected with adhesive tape.

Standard photolithography is used to pattern Au and Cu. Pleaseote that a temporary glass substrate is used in this step for easierandling of the relatively thin sample. Positive S1813

TMphotore-

ist from Microchem® is spin coated for 30 s on the Au layer andrebaked for 10 min at 90◦C. After UV illuminating, the sample iseveloped in KOH and is post-baked for 15 min at 110◦C (Fig. 3d).ubsequently, the Au layer is etched for 5 min in Aqua Regia. Thistchant is not 100% selective to Au; it also etches the Cu layer dueo the presence of the nitric acid. Cu layer is etched in Edinburgtchant for 20 min (Fig. 3e). In this step, 200 �m wide meander-haped tracks and electrical pads are defined. The glass substrate is

hen removed (Fig. 3f).

The back side of the sample is covered with PDMS (Sylgard®

84 from Dow Corning®). The PDMS is prepared with a 10:1 ratiof precursor to curing agent. It is stirred, degassed in vacuum, and

substrate. g) the back side of the sample is cast into PDMS. h) PI is dry etched usingpatterned Au and Cu as the hard mask.

is manually poured on the sample. It is partially cured in air for10 min and then in an air convection oven at 80◦C for 4 h (Fig. 3g).It results to approximately 1 mm of thickness. To enhance the adhe-sion between PI and PDMS, a thin layer of flexible glue (RaziFlexTM

from a local shop) is applied precisely behind the tracks. This stepis done manually and is not illustrated in the figure. Finally, PI isisotropically etched in a reactive ion etching (RIE) system while

Au and Cu serve as the hard mask. The etching process is done in100 sccm of O2 and 100 sccm of SF6 at 300 W with an etch rate ofapproximately 0.5 �m/min (Fig. 3h). As shown in this figure, the

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ig. 4. The camera image of the fabricated stretchable dry (a) EMG and (b) ECGlectrode is shown in this image.

dge of the Cu layer is exposed to the skin. However, the thicknessf the Cu layer is negligible compared to the surface area of the drylectrode and it did not produce any observable side-effect, evenn long-term scenarios.

Please note that the PI is patterned with precisely the sameimension of the metal layers. The width of the PI layer can beesigned more than Cu to increase the cycling reliability [32]. How-ver, it adds an extra mask in the fabrication process. It is believedhat due to the high stretchability of these tracks, it is not generallyecessary [33]. The external wires can be soldered to the electricalads. Since the Cu layer from the double-sided flex is thick enough,he wire bonding step can be done using standard soldering proce-ure. The 25 �m PI layer can favorably tolerate the relatively highemperature of soldering.

Fig. 4 shows the camera image of the fabricated dry electrodes.he EMG and ECG electrodes are shown in Fig. 4a and b, respec-ively. The yellow color of Au is visible in the figures. Fig. 5a showshe SEM image of the sample before PI dry etching. The thicknessf Cu and Au is indicated in the tilted image on the right-hand side.icroscope image of the sample on PDMS substrate is illustrated

n Fig. 5b. The width of the tracks is 200 �m. The relatively roughurface of the Au layer is due to the industrial grade of the surfacenish of the PI sheet.

. Results and discussion

.1. Mechanical characterization of stretchable electrodes

Biological organisms, including human, typically have soft

urvilinear tissues, compared to the current health monitoringquipment that is made from rigid, brittle materials. Using elas-omers with low stiffness greatly enhances the user experience.rom our previous works, it can be concluded that the pre-

d Actuators A 295 (2019) 678–686 681

sented dry electrodes will perform relatively well mechanically[33]. More precisely, the samples can withstand large mechanicaldeformations, including bending, stretching and twisting (Fig. 6).As illustrated in Fig. 6, the dry electrodes can be convenientlyelongated to more than 25%. In the next sections, the electrical per-formance of the presented dry electrodes is thoroughly discussed.

The electrical and mechanical properties of our stretchable elec-trodes depend on the design of the interconnects. Typically, instretchable interconnects, the maximum strain occurs at sharpangles. If the edges are rounded, the strain will have a more uniformdistribution. In addition, it has been shown that narrow parallelmulti-track meanders can be fabricated to decrease the strain whilemaintaining the relatively high conductivity [32]. The connectionpart of the tracks and electrodes are designed in curvilinear shapeto decrease the strain gradually. The proposed design can toleratea large range of elongations while providing desirable impedance.

3.2. Skin-electrode contact impedances

Electrical signals, which are created by muscle cells and trans-ported to the body surface via ions, i.e., Na+, Cl−, Ca+, can bemeasured by placing electrodes on the skin. The performance ofthe dry electrode is related to the signal quality of the device interms of noise and motion artifacts. Skin-electrode impedance isthe impedance between the body and the electrode which playsa major role in the quality of the received signal. A low skin-electrode impedance is generally desired. Higher skin-electrodeimpedances are associated with a lower signal to noise (SNR) ratioand increased susceptibility to artifacts and interference. Precisely,it will decrease the amplitude of the signal that is fed to the subse-quent electrical amplifier due to a loading effect, which will resultin a low SNR ratio. The major causes of noise in biopotentials aremotion artifacts and common mode interference. They are gen-erated by altering skin-electrode impedance due to movement.Having a good contact between the skin and the electrode greatlyenhances the quality of biopotential signals.

An electrical circuit model can be helpful to better understandthe skin-electrode impedance behavior. Different electrical modelswere suggested for the skin-electrode interface [38]. Fig. 7 showsa typical schematic diagram of the interface between the skin andthe electrode. The equivalent electrical circuit model for both wetand dry electrodes is also shown. The capacitance Cd representsa double layer capacitance between the skin and the electrodeduring charge transfer, while the Rd represents the leakage resis-tance. The series resistance Rg is related to the conductive gel. Dryelectrodes are characterized by the lack of a wet adhesive at theskin/electrode interface. After the electrode placement, skin startsto secrete sweat. Fortunately, it aids the electrical conduction asthe sweat is ionically conductive. It also increases the adhesion ofthe electrode to the skin due to surface tensions [12]. The conduc-tive gel in the wet electrode behaves similarly. The presence of airbubbles or gaps will lead to a capacitor Cg at the interface. Theelectrode-skin interface is modeled by a parallel connected capaci-tor Cg and resistor Rg. More detailed information concerning all thecomponents in the model is available in [39].

In this work, a two-electrode impedance measurement setuphas been used. Two electrodes are placed on the forearm at a pre-cise distance of 2 cm. The total impedance at different frequencies ismeasured using an Analog Discovery 2 impedance analyser modulefrom Digilent

TM. Please note that the measured impedance includes

body impedance as well. The frequency range is varied from 1 Hzto 10 kHz. The results are compared with the standard Ag/AgCl

wet electrodes using exactly similar setup. The measurements arerepeated ten times in each frequency. The results are illustratedin Fig. 8. It should be noted that no preparation of the skin tookplace before the measurement. The results indicate that the contact

682 P. Fayyaz Shahandashti et al. / Sensors and Actuators A 295 (2019) 678–686

Fig. 5. a) SEM images of the 200 �m wide tracks before PI dry etching. The image on the right-hand side is tilted 45◦ to better illustrate the thickness. b) Optical microscopeimage of the tracks is shown in this image. Au layer is patterned on top of PI.

trodes

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mpedance of the dry electrode is comparable to the wet electrode.he minute difference between the EMG and ECG electrodes is dueo the different surface area of the electrodes. The impedance of thery electrode changes with respect to the applied pressure as well.

transparent tape was used to attach the dry electrode to the skino, relatively moderate pressure is applied on the skin.

One of the applications of the presented technology is long termonitoring of the biopotentials. Fig. 9 illustrates the normalized

ontact impedance of the EMG and ECG electrodes at different fre-uencies. The normalization is done in order to illustrate the curves

n a single graph for comparison purposes. As expected, the con-act impedance of the dry electrode is decreased gradually withhe course of time irrespective of the frequency. The impedancesre normalized to the measured impedance in the beginning. The

is shown under various mechanical deformations.

results indicate that these electrodes can favorably be adapted inlong term applications.

3.3. Dry EMG electrode

Surface electromyography (sEMG) refers to the recording of theelectrical activity of muscles. The EMG signal can be used in manyapplications, such as the diagnosis of muscle nerve discoordination,muscle fatigue monitoring, etc. Measurement of the EMG signalis challenging, given the relatively wide frequency bandwidth of

10 Hz - 5 kHz. In addition, the received signal strongly depends onthe location of the electrodes. To collect the EMG signal, a custom-made electrical circuit transmitted the EMG signal to the PC. Inorder to limit the noise power, a bandpass filter of 50–2500 Hz

P. Fayyaz Shahandashti et al. / Sensors and Actuators A 295 (2019) 678–686 683

Fig. 7. Equivalent circuit model of the skin-electrode interface: (a) Wet and (b) Metal dry electrode.

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Fig. 9. Long-term contact impedance of dry EMG and ECG electrodes for 20 h indifferent frequencies is shown.

ig. 8. (a) Amplitude and (b) phase of the contact impedance of dry and wet elec-rodes are illustrated. Each test is repeated 10 times.

as used. The signal is amplified 800 times in two stages of 4xnd 200 × . An instrumentation amplifier (AD622, Analog Devices)mplifies the potential difference between the measurement and

he reference electrodes, to suppress common mode noise in eachnput signal. An operational amplifier (TLC2272) is also used forlter and second stage amplifier. The amplifier and filter units pro-ide tunable gain for EMG measurements. Fig. 10 shows a camera

image of EMG electrode on the forearm, the electrical circuit andthe digital oscilloscope.

Two separate tests were done for forearm and bicep. These loca-tions are chosen to ensure a high amplitude EMG signal. Separatebut similar tests are done for dry and wet electrodes. The dry andwet electrodes are placed exactly in the same position of the fore-arm and the bicep in order to be able to compare the results. Thereference electrode is placed on the bony area of the elbow. Nospecial skin preparation was performed before the test. The resultsare compared in Fig.11. The volunteer (age 26, male) was asked totighten his right hand repetitively during the test. The results indi-cate that no significant difference can be observed between the dryand the wet electrode.

The power spectral density (PSD) of the EMG signal from fore-arm is illustrated in Fig. 12. The plots show that most of the EMGpower from the electrodes were concentrated on the region below250 Hz which is expected [20]. The PSD of the dry and wet elec-trodes show a similar pattern as can be seen in the figure. A similarSNR of approximately 55 dB can be observed both in wet and dryelectrode. SNR is estimated from PSD plots by subtracting the mainsignal power which is the maximum of the PSD curve, from thenoise power which is the minimum of the PSD curve at high fre-

quencies.

684 P. Fayyaz Shahandashti et al. / Sensors and Actuators A 295 (2019) 678–686

Fig. 10. A camera image of the sEMG electrode is shown which is placed on the forearm. The electrical circuit consisting of the filter and 800x amplifier is also shown. Thesignals are sent to the PC using the digital oscilloscope shown on the right-hand side of the image.

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ig. 11. Experimental sEMG signals of a series of muscle contractions with the drynd wet Ag/AgCl are shown. The electrodes are placed on (a) forearm (b) bicep.

.4. Dry ECG electrode

The ECG signal is a vital biological potential in medical appli-ations, as this signal contains important information about theuman heart and its physical function. The design of the dry ECGlectrode is illustrated in Fig. 1b. A test scenario similar to the dryMG electrode is carried out for ECG as well. A single-lead, heartate monitor front end (AD8232 from Analog DevicesTM) is used to

cquire, amplify, and filter ECG signals. The system provides a gainf 100 in the frequency range of 0.5 Hz – 100 Hz. All ECG signals areeceived from the arm. The working electrodes are connected to the

Fig. 12. The power spectral density of EMG signals: (a) Stretchable dry electrode (b)Ag/AgCl wet electrode.

left and right arm of the volunteer (age 26, male). The Ag/AgCl ref-erence electrode is connected to the right leg thigh. Neither shavingnor skin preparation has been done before placing the electrodes.The tests for dry and wet electrodes have been done separately. In

the tests, the electrodes are placed in exactly the same positions.The results have been shown in Fig. 13. QRS complex waves, P wave,and T wave are clearly visible in the signals received by both elec-

P. Fayyaz Shahandashti et al. / Sensors an

Fig. 13. The ECG signal is received using the dry and wet Ag/AgCl electrode. Thesignals are received from the left and right arm whereas the reference electrode isconnected to the right leg thigh.

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ig. 14. The ECG signal is received in walking condition using stretchable dry andet Ag/AgCl electrodes.

rodes. No significant difference can be observed between the drynd wet electrode.

In order to investigate the sensitivity to motion artifacts due toody movements, a test scenario was designed. While the subject

s moving, ECG signals are recorded from the left and right arms.imilar to the previous test scenario, the signals were acquiredrom the stretchable dry electrode as well as the wet Ag/AgCllectrode from the same positions. The stretchable ECG electrodesere attached on the skin using a transparent tape similar to the

mpedance measurement test (also shown in Fig. 10). The electricalables were firmly secured to the boy, isolating the motion arti-acts due to cable movements. Therefore, the measured distortions mainly attributed to the skin-electrode impedance variations.ig. 14 shows that the stretchable dry electrode features relativelyess sensitivity to body motions than the conventional Ag/AgCl elec-rode. In comparison to Fig. 13, the ECG signals are received withoise and zero-line fluctuation, although the P, QRS and T wavesre clearly visible. One can observe that due to the flexibility of thery electrode, the motion artifacts is minimized.

. Conclusion

In this work, a novel flexible and wearable technology for theabrication of dry stretchable electrodes was presented. Stretch-

[

d Actuators A 295 (2019) 678–686 685

able dry electrodes were demonstrated to measure EMG and ECGsignals with relatively good performance. The proposed stretchabledry electrodes, provide on par or even more promising results com-pared to the standard wet Ag/AgCl electrode. The dry electrodeswere shown to compete equally well in terms of the received SNR.However, relatively higher insensitivity to motion artifacts can beobserved in the dry electrodes. The other advantages of the pro-posed technology include less skin irritation due to the absenceof conductive gel and no signal degradation as a result of betterskin-electrode compliance in long term tests.

The presented technology in this work fits well within thestretchable electronics field of activity. This technology features arelatively inexpensive platform for the fabrication of stretchabledry electrodes with relatively high conductivity, reliability, andnonetheless biocompatibility. Thanks to the extreme compliance,the biopotential signals can be monitored at the comfort of theuser in long-term real-world applications. This and similar worksare a stepping stone towards the realization of human biosignalsmonitoring in clinical health-care applications.

Appendix A. Supplementary data

Supplementary material related to this article can be found,in the online version, at doi:https://doi.org/10.1016/j.sna.2019.06.041.

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Electrical Engineering Department, Amirkabir University of Technology (TehranPolytechnic), Tehran, Iran. He has authored and co-authored more than 100 papersin refereed journals and international conferences. His research interests includesemiconductor devices, integrated photonics, quantum physics and quantum elec-

86 P. Fayyaz Shahandashti et al. / Sens

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Biographies

Peyman Fayyaz Shahandashti received his MSc degree in Micro and Nano Elec-tronic Devices from AmirKabir University of Technology (Tehran Polytechnic),Tehran, Iran, in 2019. His current research interests include micro/nanofabricationTechnology, Stretchable/Flexible Electronics, and MEMS.

Hamed pourkheyrollah received his MSc degree in Micro and Nano ElectronicDevices from AmirKabir University of Technology (Tehran Polytechnique) inFebruary 2019. His current research interest focuses on flexible and stretchableelectronics particularly stretchable micro-electrodes and MEMS biosensors forhealth-care applications.

Amir Jahanshahi received his PhD degree in electrical engineering from Gent Uni-versity (imec) in 2013. He is now an assistant professor in EE at AmirKabir Universityof Technology (Tehran Polytechnic), Tehran, Iran. His main research activity is instretchable / flexible electronics as well as microfluidics. His team at Micro Bio Tech-nology Lab (MBTechLab) are working towards real-world realization of polymermicro systems and biosensors interacting with the human body.

Hassan Ghafoorifard received the B.Sc. degree in Physics from Tehran Univer-sity,Tehran, Iran, in 1965, the M.Sc. and Ph.D.degrees both in Physics from Universityof Kansas, USA, in 1973 and 1976, respectively. He is currently a professor at

tronics.