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Sensors and Actuators A 206 (2014) 22– 29

Contents lists available at ScienceDirect

Sensors and Actuators A: Physical

jo u r n al homep age: www.elsev ier .com/ locate /sna

iquid silicone rubber (LSR)-based dry bioelectrodes: The effect ofurface micropillar structuring and silver coating on contactmpedance

alla Kaitainena,∗, Antti Kutvonenb, Mika Suvantob,c, Tuula T. Pakkanenb,eijo Lappalainena,c, Sami Myllymaaa,c

Department of Applied Physics, University of Eastern Finland, P.O. Box 1627, 70211 Kuopio, FinlandDepartment of Chemistry, University of Eastern Finland, P.O. Box 111, 80101 Joensuu, FinlandSIB Labs, University of Eastern Finland, P.O. Box 1627, 70211 Kuopio, Finland

r t i c l e i n f o

rticle history:eceived 3 July 2013eceived in revised form1 November 2013ccepted 18 November 2013vailable online 25 November 2013

eywords:lectrodempedanceiquid silicone rubbericropatterning

a b s t r a c t

Bioelectrode–skin contact impedances need to be minimized to enable high quality recordings of biopo-tentials. Low contact impedances can be achieved with large, effective bioelectrode–skin contact areas. Inthis study, novel dry microstructured electrodes based on electrically conductive liquid silicone rubber(LSR) were developed and characterized with electrical impedance spectroscopy (EIS) measurements.Two different micropillar structures with inter-pillar distances of 20 �m and 100 �m, respectively, werecompared with smooth LSR electrodes, as such or coated with silver (Ag). A gelatin gel based skin modelwas used in EIS measurements to overcome the problems associated with real skin measurements. Bothmicrostructurings were found to significantly decrease the contact impedance modulus. In addition, theAg coating caused highly significant (p < 0.01) decrease in impedance modulus for each surface topog-raphy, e.g. for smooth electrodes the impedance drop was roughly from 140 ± 30 k� to 50 ± 10 k� at1 Hz. The Ag coating combined with microstructuring expectedly revealed the most superior electrical

lectrical impedance spectroscopyiosignals

characteristics as the contact impedance declined even as low as 10 ± 2 k� at 1 Hz. Instead, no statisti-cally significant differences in impedance values were detected between the two microstructurings in theuncoated or Ag-coated electrodes. However, more sparse microstructuring (100) seemed to offer lowervariation, i.e. more reproducible results. In conclusion, present microstructured electrodes that do notpierce any layer of skin but improve the contact to skin might be a good alternative to conventional wet

electrodes.

. Introduction

Bioelectrodes are extensively used to detect electric signalsrising from the inside of the body due to the electrochemi-al activity of a certain class of cells, known as excitable cells.ioelectric phenomena, such as electrocardiography (ECG), elec-roencephalography (EEG) and electromyography (EMG), can beecorded from the surface of the body. These biopotentials areecorded using bioelectrodes that convert the ionic currents fromhe body to electric currents [1].

There is a wide variety of different types and shapes of metallectrodes, such as plate and cup electrodes, suction electrodes and

oating electrodes [1]. Most widely used metal electrode type ishe Ag/AgCl electrode. One major challenge in recording biopoten-ials is the electrically insulating outer layer of the skin, i.e. stratum

∗ Corresponding author. Tel.: +358 50 4138227.E-mail address: salla.kaitainen@uef.fi (S. Kaitainen).

924-4247/$ – see front matter © 2013 Elsevier B.V. All rights reserved.ttp://dx.doi.org/10.1016/j.sna.2013.11.020

© 2013 Elsevier B.V. All rights reserved.

corneum. To improve the conductivity of the electrode–skininterface, skin abrasion and/or a conductive gel or paste isconventionally utilized with metal electrodes. Both of these prepa-ration techniques subsequently decrease electrode–skin contactimpedance allowing the measurements with a sufficient signal-to-noise ratio. However, these preparations are time-consuming(i.e. long application time, stabilization time and cleansing timeafter use), and may cause discomfort or skin damage to thepatient. Furthermore, smearing of the electrolyte gel may leadto short-circuiting when electrodes are placed close to eachother. In long-term measurements, gel drying is one major prob-lem.

To overcome the difficulties mentioned above with metal elec-trodes, other kinds of electrodes have been introduced. One of theseinnovations is the hydrogel based electrode (e.g. [2,3]). Hydrogelelectrodes do not need gel application but usually they require

skin preparation. They are flexible but normally they cannot bereattached to the skin if they are misplaced or removed. Further-more, hydrogel electrodes do not suit for performing measurement

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S. Kaitainen et al. / Sensors

n hairy areas and they are somewhat sensitive to gel drying inong-term measurements, too.

Recent research to eliminate the need of the skin abrasion andvoid the electrolyte gel drying has led to several new dry electrodepproaches [4–10]. Some of the electrodes even have microneedleso better keep them in place and to reduce skin contact impedancey penetrating through the stratum corneum [11–16]. It has beenhown that dry electrodes perform at least as well as the conven-ional Ag/AgCl electrodes [4–7,17,18]. In long term measurementshe dry electrodes can perform even better than wet electrodes [5].

One major disadvantage of the most widely used solid elec-rodes (e.g. metal plate or cup) is their inability to conformo the irregularly shaped skin surface that additionally canhange its local curvature during movement. To avoid such prob-ems, different types of flexible electrodes have been developed.

idely used materials for flexible electrodes are conductiveolymers such as poly(dimethytsiloxane) (PDMS) and poly(3,4-thylenedioxythiophene) (PEDOT) [4,5,18–22] but other materialsre used, too [6,7]. Conductive polymers are popular because theyre inexpensive, biocompatible and they can be micro-molded tochieve a desired shape.

Recently, a new method has been introduced to create poly-er surfaces with accurately defined micro- and nanoscale features

23]. Previously, we applied this methodology for electricallynsulating PDMS to produce microstructured disks, which wereubsequently coated with thin films of different electrode mate-ials [19]. In this study, the method was further developed to thelectrically conductive liquid silicone rubber (LSR). The aim was tolarify whether surface micropillar structuring with or without aroper coating can be used to provide improvements for electrodeerformance and functionality allowing their utilization as biopo-ential electrodes without gel. We varied the density of micropillartructure and developed a proper skin model for preliminary test-ng. Our micropillars do not pierce any layer of skin but they merelyncrease the effective contact area and allow easier flow of moisturend air at the skin–electrode interface.

. Materials and methods

.1. Preparation of LSR sample disks

A micro mold insert applied in the injection molding of LSR sam-les was fabricated by structuring electropolished aluminum foil

cm × 3.5 cm (0.25 mm thick, 99.997% Al, Puratronic®, Alfa Aesar)ith a microworking robot (RP-1AH, Mitsubishi Electric). A detailedescription of the pretreatment of the aluminum foil and microtructuring of the mold insert has been given elsewhere [24,25].

The size and shape of microdepressions on the aluminum foilas controlled with a micro working needle made of tungsten

arbide. In this work the cylindrical needle tip with a diameterf 60 �m was used. The depth and distance of the microdepres-ions were adjusted by parameters of the robot program. The size ofhe microstructured areas of aluminum foil was 4 mm × 4 mm. The

icrodepressions possessing an ordering of cubic symmetry hadhe diameter and depth of 80 �m. Two interdepression distancesf 20 or 100 �m were applied resulting in a total of 1600 or 484icrodepressions, respectively, in the microstructured area. The

tructured aluminum foil was supported by gluing it onto a steellate (thickness 0.5 mm) with a thermostable epoxy resin (Loctiteysol 9492 A&B). At the final stage the mold insert was cut into cir-ular shape 25 mm in diameter. The circular mold insert contained

everal separate microstructured areas.

LSR sample disks with microstructured areas were preparedith a DSM Midi 2000 extruder and microinjection-moldingachine. Prior to extrusion, A and B components of Silopren

tuators A 206 (2014) 22– 29 23

LSR 2345/05 kit (Momentive performance materials, Leverkusen,Germany) were mixed manually in the ratio of 1:1. The followingprocessing parameters were used in replication of the mold struc-tures in LSR elastomer: the rotation speed of screws 50 rpm, screwtemperature 25 ◦C, mold temperature 165 ◦C and air pressure ofinjection piston 5 bars. The cross linking time (dwell time) of theelastomer in the mold was 10 s.

2.2. Deposition of Ag coating on LSR sample disks

A part of the LSR sample disks were coated with silver (Ag).Before the coating deposition sample surfaces were cleaned in vac-uum with high purity argon (99.999% pure Ar, Instrument Argon5.0, Oy AGA Ab, Espoo, Finland) sputtering for 4 × 40 s, at 30 mAcurrent, ̨ ≈ 45◦ angle. Then the sample disks were first coatedwith a very thin film of titanium (Ti) using magnetron sputter-ing technique (Stiletto Serie ST20, AJA International Inc., NorthScituate, MA, USA) to improve the attachment of the Ag coat-ing. Finally, they were coated with about 500 nm thick Ag layerusing magnetron sputtering (3 × 3 min, 480 V, ̨ ≈ 45◦, d = 30 cm).The examination of the uncoated and Ag-coated microstructuredsamples was undertaken with scanning electron microscope XLESEM TMP (FEI Company/Oy Philips Ab, Brno, Czech Republic).

2.3. Preparation of LSR electrodes for impedance testing

After the depositions, the 4 mm × 4 mm sized microstructuredareas were cut out from the sample disks with a special custom-made tool. In addition, smooth areas were also cut out fromthe sample disks for comparison purposes. The pieces werethen glued to 6 mm × 8 mm “999 fine” silver pieces (thickness0.25 mm, purchased from Hopeasavi, Helsinki, Finland) with sil-ver epoxy (EPO-TEK H20E, Epoxy Technology, Billerica, MA, USA).Then, Kapton® tape insulator was attached around the electrodes(Fig. 1b). Altogether, we had a total of 20 samples, 3–4 electrodesof each type. Below, smooth electrodes will be called smooth andthe microstructured ones 20 and 100. Ag-coated electrodes will beidentified with smooth c, 20 c and 100 c.

2.4. Skin model

We used a 28% (w/w) gelatin gel as a skin model in electricalimpedance spectroscopy (EIS) measurements. The gelatin powder(Dr. Oetker Suomi Oy, Helsinki, Finland) was made of pig skin. Thepowder was dissolved in physiological saline solution (0.9% NaCl)in hot water bath. The solution was poured in a Petri dish and left toset in the fridge for a few hours before the EIS measurements. Thethickness of the gelatin gel was 5 mm. Prior to the measurements,the gelatin gel was left to stabilize at room temperature in order toeliminate possible changes in gel’s mechanical and electrical prop-erties due to changes in gel’s temperature. For the measurementsof each electrode, a piece of gel (14 mm in diameter) was cut fromthe gel in the Petri dish.

2.5. EIS measurements

The EIS measurements were conducted using a Solartron 1260impedance/gain-phase analyzer coupled to a Solartron 1287 elec-trochemical interface (Solartron Analytical, Farnborough, UK). Themeasurements were carried out in the frequency range from 0.1 Hzto 1 MHz by applying a sinusoidal alternating excitation voltage of100 mV without any DC offset. Data acquisition and data analysis

were performed using software packages ZPlot and ZView (ScribnerAssociates, Inc., Southern Pines, NC, USA).

The measurement configuration is presented in Fig. 1. A custom-made measurement apparatus (introduced in [26]) was used to

24 S. Kaitainen et al. / Sensors and Actuators A 206 (2014) 22– 29

Fig. 1. (A) Setup used in EIS measurements, (B) electrodes ready for measurements (shown in centimeter scale). From left to right: smooth, 100 c and underside of anelectrode showing a larger area fine silver piece glued to the LSR electrode. The round-shaped insulator around the electrode is Kapton® tape. (C) Schematic representationof the setup. A small weight was applied to ensure constant contact pressure between the electrode and gelatin gel based skin model.

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modulus of gelatin gel was much lower than that of the samples(Fig. 3). In addition, the variations in impedance moduli between gelbatches were small (at 1 Hz, RSD = 20.0%; at 1000 Hz, RSD = 10.8%;

Fig. 2. SEM images of Ag-coated microstructure

ompress the piece of gelatin gel and the electrode in contact with constant force. There was 1 min stabilization time after placinghe electrode and the gel in the measurement apparatus prior totarting the measurement scan. The Kapton® tape in the electrodesrevented short-circuiting between the stainless steel rods. Eachlectrode was measured twice without changing the gel betweenhose two measurements.

.6. Statistical analyses

The results are expressed as mean ± standard deviation (SD).PSS software (IBM SPSS Statistics, version 19.0) was used toetermine the statistical significance of the observed differencesetween the samples. Mann–Whitney U test was used, and sig-ificance level p < 0.05 was considered as significant and p < 0.01s highly significant. Statistical differences in impedance moduliere studied at 1 Hz, 10 Hz, 100 Hz and 1 kHz to cover a relevant

requency range of ECG, EEG, EMG, etc.

. Results

The preparation of LSR sample disks succeeded with optimizednjection molding process parameters. Microscopic examinationevealed that both micropillar structurings were very homoge-

eous throughout the samples without any artifacts (Fig. 2). Theone-shaped pillars turned out to be advantageous in sputteringrocess allowing the formation of continuous Ag layer (conformaloverage) deposited on the vertical edges of pillars, too. The Ag

samples. (A) Sample 20 c and (B) sample 100 c.

coating adhered well to the LSR disks without any signs of cracksor delamination during the EIS measurements.

The used measurement setup was reliable enabling the col-lection of the EIS data with high reproducibility. The impedance

Fig. 3. Impedance modulus for pure gelatin gel and 20 c and 100 with gelatin gel.Samples from three different gelatin gel batches were used in these measurements.

S. Kaitainen et al. / Sensors and Ac

Table 1Contact impedances (mean ± std) of all electrodes.

1 Hz |Z| (k�) 10 Hz |Z| (k�)

Type Uncoated Ag-coated Type Uncoated Ag-coated

Smooth 135.0 ± 25.8 46.3 ± 7.4 Smooth 135.0 ± 24.4 33.2 ± 6.020 95.6 ± 33.9 9.0 ± 1.4 20 88.0 ± 30.1 3.8 ± 1.2100 89.1 ± 17.8 10.7 ± 0.9 100 82.1 ± 15.4 4.5 ± 0.7

100 Hz |Z| (k�) 1000 Hz |Z| (k�)

Type Uncoated Ag-coated Type Uncoated Ag-coated

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Smooth 46.9 ± 10.9 25.6 ± 5.1 Smooth 8.7 ± 1.3 12.8 ± 1.020 28.6 ± 9.0 2.93 ± 1.3 20 5.9 ± 1.7 2.5 ± 1.0100 30.1 ± 7.3 3.22 ± 0.7 100 7.5 ± 1.8 2.7 ± 0.4

= 19) and thus electrode types could be compared reliably becausehe relevant differences between electrode types are larger.

According to Fig. 4 in which the Bode plots for all electrodeypes are presented and Table 1 where the numerical mean val-es are given, it is clear that the Ag coating significantly decreaseshe impedance modulus at low frequency region (<1 kHz). Thisecrease was more pronounced in microstructured electrodesFig. 4b and c) compared to smooth ones (Fig. 4a). At higher frequen-ies (>1 kHz), surprisingly the impedance modulus for uncoatedlectrodes was lower than for coated ones. The phase data showhat the uncoated and the coated electrodes had their own kindf polarization behavior. Only at very low frequencies (<1 Hz), theolarization was greater in coated than in uncoated electrodes. Athe most relevant biosignal region (1 Hz–10 kHz), the polarizationas minor in coated electrodes.

Statistical analysis of the EIS data performed at four discreterequencies (1 Hz, 10 Hz, 100 Hz and 1 kHz) demonstrated that Agoating leads to a highly significant improvement in the contactmpedance modulus at all analyzed frequencies in all electrodeypes (Fig. 5), as was expected according to previous results [19].

Comparison between uncoated electrodes, i.e. smooth versusicrostructured electrodes (20 and 100) revealed that there was at

east significant improvement in the impedance modulus in almostvery case (Fig. 5). In general, standard deviations were clearlyarger in the case of uncoated electrodes. No statistically significantifferences were detected between the two microstructures.

When comparing the Ag-coated samples a highly significantifference in impedance modulus was detected between smoothamples and microstructured ones at all frequencies (Fig. 5). Notatistically significant differences were detected between the twoicrostructures in coated samples, either.

. Discussion

This study was carried out to determine whether LSR electrodeerformance can be enhanced with microstructuring and coating as

nvestigated using a skin model. Carbon-loaded LSR electrodes areommonly used in stimulation purposes, such as in transcutaneouslectrical nerve stimulation (TENS), but LSR’s high impedance isestricted its utilization in recording applications. However, opti-ized surface modifications as successfully shown in this study,

an offer new encouraging possibilities. In order to keep this studyell focused we chose to include here the results from the skinodel measurements only. Consequently, the focus of our subse-

uent paper will be on human skin impedance measurements andiopotential (especially EEG, ECG and EMG) recording experiments.

The bioelectrode–skin contact impedance plays a crucial role

n the performance of the electrode. The contact impedances needo be minimized to enable high quality recording of biopotentialshroughout the testing. If the contact impedance is too high theignal-to-noise ratio is decreased leading to poor data [27]. One

tuators A 206 (2014) 22– 29 25

important factor regarding contact impedance is the contact areaof the skin–electrode interface. The larger the contact area thesmaller the contact impedance [6,28,29]. Using microstructuringthe effective area of the electrode can be increased without mak-ing the electrodes very large. The microstructures in this studyprovided 61% (microstructure 100) and 200% (microstructure 20)increase to the total calculated electrode area. Both microstruc-turings clearly improved the electrodes’ performance except 20at 1 Hz (Fig. 5). When a metallic coating is added the improve-ment is even more evident. However, no statistically significantdifferences were detected in impedance modulus between the twomicrostructures when comparing the uncoated or coated samples.Therefore the improvements achieved cannot simply be explainedby the increase of the active surface area due to the microstructur-ing which could theoretically give a maximum of 200% increase.However, this cannot be achieved especially at 1 Hz with uncoatedelectrodes. Therefore, the measured improvements are also due tothe improved contact between the electrode and the skin phantom.The flow of moisture between the microstructuring also decreasesthe contact impedance. In real skin measurements the flow of mois-ture is really beneficial since sweat can easily spread between themicrostructuring to cover the whole electrode area and conse-quently decrease the contact impedance.

In this study, the most significant improvements in con-tact impedance modulus were achieved by a combination ofmicrostructuring and Ag coating although even the microstruc-turing or Ag coating alone improved contact impedance modulus.Furthermore Ag coating improved the polarization behavior asexpected. At the relevant frequency range (1 Hz–10 kHz) the Ag-coated electrodes are less polarizable than the uncoated ones(Fig. 4). That characteristic is essential when recording small biosig-nals because the measured signals will be only slightly attenuatedand distorted [1].

Recently, several new dry electrode designs have beenintroduced. These dry electrode solutions can be roughly dividedinto two categories: skin surface electrodes [4–10] and skinpiercing electrodes [11–16]. Our LSR-electrodes fall into the firstcategory. The electrodes that do not pierce the outer layer of theskin rely on different materials (e.g. [5,6]) or surface texturing (e.g.[9,10]). Our solution to use polymer as the electrode material andsurface texturing and coating as a way of improving the contactimpedance is not a novel solution as such, but combined with themanufacturing techniques our solution can provide an economicalalternative. The motivation for studying dry electrodes and theirimprovement is based on their many merits, such as rapid andeasy attachment on skin, stable electrode performance through-out the long-term measurements and user comfort. However, theproblems associated with dry electrodes are related to e.g. highelectrode–skin contact impedance and susceptibility to motionartifacts. As Meziane et al. [30] concluded, dry electrodes could bewidely used if these problems could be overcome. Dry electrodeshave many application areas in clinical medicine – they can beused in for example ambulatory monitoring of EEG/ECG [21,31] andrehabilitation [18]. Other application areas are for example healthmonitoring during normal daily life [20] and sports [4].

Comparison of our results to those found in the literature is notstraightforward due to the great variability in impedance measure-ment setups. In our measurement configuration (Fig. 1) we onlymeasured one electrode at a time with skin model using a largearea SS rod as a counter electrode, whereas many papers reportvalues from human skin measurements in which two identical testelectrodes were attached either unprepared or prepared skin. Also,

the contact impedance values from different measurement setupsdiffer a lot from one to another. For example, Baek et al. [5] reportedvalues under 2100 k� for their PDMS electrodes and values under500 k� for standard Ag/AgCl electrodes on prepared skin at the

26 S. Kaitainen et al. / Sensors and Actuators A 206 (2014) 22– 29

Fig. 4. Bode plots of EIS data for all sample types (n = 3 or 4). Each electrode was measured twice. Contact impedance moduli are shown on the left column and phase angleso c.

frsqisiqeattrkiitumIoao

n the right column. (A) Smooth and smooth c, (B) 20 and 20 c and (C) 100 and 100

requency range of 1–1000 Hz. On the other hand, Lin et al. [6]eported values under 10 k� for their own foam electrode andtandard Ag/AgCl electrodes on prepared skin at the same fre-uency range. Searle and Kirkup [32] have measured the contact

mpedance of Ag/AgCl electrodes to be 150 k� on unpreparedkin at 57 Hz, whereas Fiedler et al. [9] reported their Ag/AgClmpedance values to be under 3 k� on prepared skin at the fre-uency range of 1–1000 Hz. At the mentioned frequency range ourlectrodes have the contact impedance under 130 k� (uncoated)nd under 47 k� (Ag-coated). For Ag-coated microstructured elec-rodes the contact impedance stays under 10 k�. Another problemo make straight comparisons with other published studies iselated to the rather small size of our electrodes. It is wellnown that the contact impedance can be reduced by increas-ng the contact area between skin and electrode [6,28,29]. As thempedance is proportional to the electrode area, we estimatedhat the contact impedance values of our electrodes might bender 30 k� (uncoated) and under 10 k� (Ag-coated) at the afore-entioned frequency range if they had a diameter of 10 mm.

t indicates that our microstructured LSR electrodes, even with-ut coating, might have appropriate electrical properties if theyre scaled to better suit skin measurements. The performancef the LSR electrodes might be further improved if a more

conductive LSR is used as can be expected according to the previousstudies [10,33].

It has been recommended by EEG device manufacturers that theskin–electrode impedance in EEG recordings need to be achievedbelow 5 k� enabling high-quality recordings without attenua-tion in signal amplitude and signal-to-noise ratio. However, withmodern high input-impedance amplifiers, the attenuation of thesignal is not a major concern anymore. Ferree et al. [34] showedthat there was no significant change in EEG amplitude as scalp-electrode impedance increased from less than 5 k� (abraded skin)to 40 k� (intact skin) and they concluded that high-quality EEGcan be recorded without skin lesions. Most commercial biosignalamplifiers have in-built impedance checking circuit that normallyoperates about at 10 Hz frequency. In this study, the impedancesdetermined at 10 Hz varied between 140 k� (smooth) and 3 k�(20 c).

As expected at frequencies high enough, smooth andmicrostructured electrodes behave in a similar way with andwithout coating. However, the frequency where the curves

merge moves toward higher frequencies as the active surfacearea increases (Fig. 4). At very high frequencies, the impedancemodulus of uncoated electrodes was even lower than that of thecoated ones. This behavior was the most distinct in electrodes

S. Kaitainen et al. / Sensors and Actuators A 206 (2014) 22– 29 27

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ig. 5. Mean values and standard deviations of impedance moduli of all electrode

p < 0.05), and highly significant differences are marked with ** (p < 0.01). On the tohe effect of microstructuring is shown.

aving inter-pillar distance of 100 �m (Fig. 4). The phenomenonan be explained on the basis of the Maxwell–Wagner interfacialolarization effect exhibiting anomalous dispersions in dielectricaterials comprising of separate layers [35]. Therefore, the coated

lectrodes used in this study were more prone to interfacial polar-zation effects at high frequencies (>10 kHz) since they consistf superimposed LSR, Ti and Ag layers. However, the commonlyeasured biosignals are in the frequency range of 1–10 000 Hz.

or example, EEG signals are in the range of 0–100 Hz, EMG signalsre 0–10 000 Hz and EEG 0–100 Hz [1]. At that frequency range ourg-coated microstructured electrodes performed well (Fig. 4).

Agar gel phantoms have been often used in bioelectrode stud-es [36,37]. In our study, instead of agar we chose gelatin made oforcine skin because it is comparable to human skin in its struc-

ural and biochemical properties [38]. The properties of human skinary with time and from site to site [39] and using a gel made ofelatin ensured that the properties of the gel remained stable fromeasurement to measurement and the results were repeatable. We

at four discrete frequencies. Statistically significant differences are marked with *, the effect of Ag coating on each electrode type is illustrated. On the bottom row,

used 0.9% NaCl solution as a solvent which ensured an appropriateconductivity of gelatin gel mimicking the conductivity of the realhuman skin. The gel strong enough to withstand the pressure dur-ing the measurements without tearing was accomplished with the28% (w/w) solution [40]. We observed that a higher mass fractionmight lead to incomplete dissolution of the gel powder and the gelwould not be homogeneous.

Variations in the dimensions of the electrodes and gelatin gelcaused slight differences in impedance spectra of similar type ofelectrodes (Fig. 4). Actually we measured the total impedance of thegel, LSR sample, silver epoxy, “999 fine” silver piece, contacts andleads. The differences can mainly be explained with small changesin the thicknesses of gelatin gel pieces and LSR samples. Neverthe-less, the repeated measurements of each sample gave consistent

results. In Fig. 5, it can be seen that the standard deviations for100 are smaller than for 20 or smooth. The reason to this mightbe that the denser pillar structure in 20 does not allow the skinphantom to conform with the pillars as well as in the case of 100.

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n 100 the point contact pressure is higher than in 20 or smoothhich improves the electrode’s contact to skin model because theicropillar sides are utilized, too. That makes the measurementsith electrode 100 more reproducible.

Our future studies will be focusing on the realization of real elec-rode sets for EEG, ECG and EMG studies. Further investigationsre necessary to test different micropillar configurations, sizes oflectrodes, their long-term behavior and durability of coatings.

. Conclusions

We successfully developed and tested different kinds of flexiblery electrodes with micropillar structurings and Ag coating for thessessment of their electrical properties. Decrease of 40 k� (30%)n the contact impedance modulus was achieved with microstruc-uring and a decrease of 80 k� (90%) was achieved with the Agoating at 1 Hz. The skin phantom studies gave clear evidences thaticrostuctured LSR electrodes have a potential to be used as skin

lectrodes if scaled to a larger size. We conclude that our designs not significantly different from other dry electrode designs buthat there may be advantages in terms of cost since this solution iseadily feasible in to mass production.

cknowledgements

We thank Dr. Tiina Rasilainen for sharing her expertise in surfacetructuring and MEng Hannu Korhonen for help with the depo-itions. The work was supported by the strategic funding of theniversity of Eastern Finland (NAMBER project).

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Biographies

Salla Kaitainen received her MSc degree in medical physics from the Departmentof Applied Physics, University of Eastern Finland, Kuopio, Finland in 2010. She iscurrently working toward the PhD degree. Her interests lie in nano and micro scalesurface modification of biomaterials and their applications in cell culture controllingand biosignal monitoring.

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ntti Kutvonen graduated from the University of Eastern Finland with an MScegree in Materials Chemistry in 2010. His Master of Science thesis concerned theurface chemistry, microstructures and molding techniques of silicon elastomers.utvonen is currently working with petrochemicals as a laboratory chemist in theGS. His work interests are in gas chromatography and chemical engineering.

ika Suvanto is currently the research director and head of scientific materialesearch and infrastructure unit SIB Labs at the University of Eastern Finland. Prioro join to SIB Labs he served as the director of the Special Materials Research Centeror nine years. He received his PhD degree in physical chemistry in 1999 from theniversity of Joensuu. His research interests lie in surface and coating technology.urrently he is actively involved in researches related to development and char-cterization of functional polymer surfaces, composites, and coating applicationsased on nano and micro scale features and additives.

uula Pakkanen has been a professor of Materials Chemistry at University of Joen-uu 1994–2010 and at University of Eastern Finland since 2010. She received herSc degree in Chemical Engineering from Helsinki University of Technology in

973 and PhD degree in Chemistry from State University of New York at Stonyrook in 1978. Her current interests include organometallic chemistry, coordination

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polymerization catalysts, polymer nanocomposites and self-assembly of plasmonicnanoparticles. She has published more than 150 scientific articles.

Reijo Lappalainen has been a professor of Biomaterials Technology at the Univer-sity of Kuopio 1999–2010 and at the University of Eastern Finland since 2010. Hereceived his MSc degree in Physics from Helsinki University in 1982 and PhD degreein Physics in 1986. Recently R. Lappalainen has been very active in the develop-ment of ultra-short pulsed laser ablation techniques, wood distillate coatings andnovel nanomaterials. Furthermore, new methods related to spectroscopic analysisand lithography have been developed and utilized. He has published more than 160peer reviewed scientific articles and more than 30 patents.

Sami Myllymaa received his MSc degree in Electrical Engineering in TampereUniversity of Technology, Finland in 2001, and the PhD degree in Biomedical Engi-neering from University of Eastern Finland, Kuopio, Finland in 2010. Currently

he is a post-doctoral researcher in the Department of Applied Physics at theUniversity of Eastern Finland. His research interests lie at the junction betweenphysics, medicine and biological sciences, especially focusing on the develop-ment of novel measurement techniques for biosignal acquisition and saliva-baseddiagnostics.