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Medical Engineering & Physics 33 (2011) 967– 972

Contents lists available at ScienceDirect

Medical Engineering & Physics

jou rna l h omepa g e: www.elsev ier .com/ locate /medengphy

he use of hydrogel as an electrode–skin interface for electrode array FESpplications

len Coopera,∗, Anthony T. Barkerb, Ben W. Hellerc, Tim Goodb, Laurence P.J. Kenneya, David Howardd

Centre for Health, Sport & Rehabilitation Science Research, University of Salford, UKDepartment of Medical Physics and Clinical Engineering, Sheffield Teaching Hospitals NHS Foundation Trust, UKCentre for Sports Engineering Research, Sheffield Hallam University, UKSchool of Computing, Science & Engineering, University of Salford, UK

r t i c l e i n f o

rticle history:eceived 22 October 2010eceived in revised form 16 March 2011ccepted 16 March 2011

eywords:lectrode arraysunctional electrical stimulation

a b s t r a c t

Functional electrical stimulation is commonly used to restore function in post-stroke patients in upperand lower limb applications. Location of the electrodes can be a problem hence some research groupshave begun to experiment with electrode arrays. Electrode arrays are interfaced with a thin continuoushydrogel sheet which is high resistivity to reduce transverse currents between electrodes in the array.Research using electrode arrays has all been conducted in a laboratory environment over short timeperiods but it is suspected that this approach will not be feasible over longer time periods due to changesin hydrogel resistivity.

lectrical FE modellingydrogellectrode resistivitylectrical stimulation efficiencylectrical stimulation focality

High resistivity hydrogel samples were tested by leaving them in contact with the skin over a sevenday period. The samples became extremely conductive with resistivities reaching around 10–50 � m. Theeffect of these resistivity changes was studied using finite element analysis to solve for the stationarycurrent quasi-static electric field gradient in the tissue. Electrical stimulation efficiency and focality werecalculated for both a high and low resistivity electrode–skin interface layer at different tissue depths. Theresults showed that low resistivity hydrogel produced significant decreases in stimulation efficiency and

resis

focality compared to high

. Introduction

It is now well established that, for subjects with foot drop as result of stroke, functional electrical stimulation of the peronealerve during the swing phase of gait results in significant improve-ents in walking [1–3]. However, a survey of foot drop stimulator

sers showed that difficulty in locating the correct electrode site ishe principal reason (after improvement in mobility) for discon-inuing FES use [4]. Although training subjects in how to placelectrodes correctly is believed to reduce the incidence of suchroblems, anecdotal evidence suggests that correctly locating thelectrodes on a daily basis remains a significant challenge for many.he problem of electrode location in the literature mainly focuses

n the lower limb but similar problems are also present in upperimb applications [5].

Over recent years, several groups have demonstrated theotential for using an electrode array in combination with a pro-

∗ Corresponding author at: School of Engineering, Manchester Metropolitan Uni-ersity, John Dalton Building, Chester Street, Manchester M1 5GD, UK.el.: +44 161 2471628.

E-mail address: g.cooper@mmu.ac.uk (G. Cooper).

350-4533/$ – see front matter © 2011 IPEM. Published by Elsevier Ltd. All rights reserveoi:10.1016/j.medengphy.2011.03.008

tivity hydrogel.© 2011 IPEM. Published by Elsevier Ltd. All rights reserved.

grammable multi-channel stimulator to enable software-controlof stimulation for both upper and lower limb applications [6–10].Such approaches use software to move an active area (virtual elec-trode or VE) around the array and, by changing the number/locationof activated electrodes, may also adjust the size and shape of theactive area.

One key problem in realising such an approach is interfacing theelectrode array with the skin. In conventional surface FES systems,the electrode is normally formed from a metal mesh backing, uponwhich is laid a thin (typically one or two mm thick) highly con-ductive and adhesive hydrogel layer. This hydrogel layer conductscurrent from the backing to the skin, acts to hydrate the skin andmaintains good mechanical skin–electrode contact.

There are clear benefits associated with the use of highly con-ductive hydrogel as an interface layer for conventional, discreteelectrodes. However, in the case of electrode arrays, a highlyconductive continuous interface layer would lead to significanttransverse currents within the hydrogel, thus negating the selec-

tion of a focussed active region on the array. For a particular arraygeometry, an increase in transverse currents in the hydrogel layeris associated with a reduction in current density in the tissue and aconsequent increase in the electrode current required to stimulatea given nerve [9–11]. To avoid this effect, Sha [10] recommended

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968 G. Cooper et al. / Medical Engineering & Physics 33 (2011) 967– 972

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ig. 1. Hydrogel samples stuck to the skin. (a) Samples were placed in the approlastowrap bandage.

hydrogel resistivity in excess of 500 � m, in combination with a mm intra-electrode gap size for a lower limb application. Kuhnt al. [11] used a cylindrical model of the forearm and showedhat, in order for the effect of stimulation with an array to beomparable with that of stimulation with an isolated electrode,ydrogel resistivities should range from 1500 to 4500 � m, depen-ent upon the inter-electrode array gap size (1–3 mm). A study byha et al. also demonstrated a secondary benefit, namely a reduc-ion in stimulation-induced discomfort, associated with the use of

high resistivity (25 k� m) interface layer [12].Whilst the timescales of these experiments is not usually

uoted, it is believed that most, if not all of them have taken placever a few hours or less. As the electrical properties of the hydro-el layer have a major impact on the extent of current spread andence on efficiency and focality, a practical electrode array and itsydrogel should ideally have electrical properties that are stablever periods of weeks, or longer.

This paper reports a study on the effects of prolonged skin con-act of one normal volunteer on the resistivity of hydrogel. Thisffect was investigated experimentally by measuring the change inesistivity with skin contact time in four different hydrogel samples.hese results were then used in a finite element model to predicthe effects on stimulation efficiency and focality.

. Experimental investigation of the effects of prolongedydrogel skin contact

.1. Methods

Following ethical approval (University of Salford ethics commit-ee REPN09/124), a single subject was recruited to the study. Theubject was male, 32 years old and had no skin conditions or health

roblems.

Four different hydrogel types with varying resistivities wereested: SCBZAB-05M, SCBZAZ-05S, PR90063 and AG3AM03-10W05 (Sekisui Plastics Company Ltd, Osaka, Japan). Each sampleas cut to approximately 31 mm × 31 mm.

able 1nitial resistivity and thickness of the hydrogel samples used in the experiment.

Sample 1: hydrogeltype SCBZAB-05M

Sample 2: hydrtype SCBZAZ-05

Initial resistivity (� m) 1230 485

Thickness (mm) 0.5 0.6

te intended location of the electrode array. (b) Samples were then covered with

The electrical properties of each sample were measured as fol-lows. Lead foil was fixed to one side of the sample and the other sidewas adhered to a copper plate. The length and width of the sam-ples were measured using a ruler and the thickness was measuredusing a Vernier gauge micrometer. A 1 kHz sine wave generator(5 V peak to peak output) was applied across the sample in serieswith a resistor, R2 (whose value was chosen to be similar to that ofthe sample to ensure). Voltages were measured across the circuit(sample plus resistor in series, V1) and across the resistor alone (V2).Phase shifts between the two measured voltages were also mea-sured to enable the sample to be modelled as a parallel resistor andcapacitor, however phase shifts were found to be negligible or verysmall in all cases showing the samples to be almost entirely restive.All electrical measurements were made using a digital oscilloscope(Tektronix THS720A). The resistivity of the sample, �S was calcu-lated using Eqs. (1) and (2) where A was the area of the sample andd was the thickness of the sample (shown in Table 1).

XS ≈ R2 ·[

V1

V2− 1

](1)

�S = XsA

d(2)

Initial measured resistivities and thicknesses are shown in Table 1.At the start of the first day each of the four sample’s electri-

cal properties and dimensions were measured, using the methoddescribed above. Each sample was then stuck to the skin of the testsubject on the lateral aspect of the lower leg (Fig. 1a). In this studythe samples were lead foil backed and covered by an elastowrapbandage (Fig. 1b) to simulate clinical application.

The samples were removed after being worn for 6 h and theresistivity of each sample measured, using the technique described

above. The 6 h period of wear was chosen to represent the timean FES user might have electrodes on their skin each day (it isappreciated that this time period will vary dependent upon theuser). The samples were placed in an air-tight bag overnight andthe procedure was repeated for seven consecutive days.

ogelS

Sample 3 hydrogeltype PR90063

Sample 4: hydrogel typeAG3AM03M-P10W05

54.9 302000.6 0.3

G. Cooper et al. / Medical Engineering & Physics 33 (2011) 967– 972 969

0 1 2 3 4 5 6 710

1

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time, days

resis

tivity,

ohm

.mscbzab-05, thickness 0.5mm

scbzaz-05s, thickness 0.6mm

pr90063, thickness 0.6mm

ag3am03, thickness 0.3mm

Fig. 2. Graph showing the change in resistivity of hydrogel with time when in con-td

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Fig. 4. Stimulation efficiency calculated from the maximum electric field gradientdivided by the total stimulation current applied equally between all electrodes andderived from the SEMCAD finite element model. Efficiency percentage is calculatedfrom the maximum stimulating function per unit current (electric field gradient inx direction per unit current applied) normalised by the value of the most efficient

tivity of the hydrogel interface layer from high to low, 1000 � mand 50 � m respectively, to approximate the changes to the highresistivity hydrogel after being in contact with the skin for a 7 dayperiod (as demonstrated in Section 2).

act with one healthy test subject’s skin (measurements taken at the end of theay).

.2. Experimental results

Fig. 2 shows the resistivity of the 4 different hydrogel sampleslotted against time. The electrical resistivity of all 4 hydrogel sam-les was observed to drop dramatically over a seven day period,o levels of between 10 and 50 � m. The rate of resistivity reduc-ion is greatest for the sample with the highest initial resistivity.lthough the samples are not all the same thickness and the exper-

ment was conducted on only one normal volunteer the results given indication of the resistivity reduction that could be expectedn high resistivity hydrogel electrode arrays used in surface FESpplications.

. Modelling of the electrode array, interface and tissue

.1. Finite element model

The low frequency solver of SEMCAD (Schmid and Partnerngineering AG, Zürich, Switzerland) was used to model a metal4 electrode array (8 × 8 format), with each electrode being

mm × 8 mm and with an inter-electrode gap of 3 mm. The arrayodelled is similar to that used in the ShefStim trial which used an

ntelligent 64-channel stimulator combined with an electrode arrayo correct drop foot in post-stroke and Multiple Sclerosis sufferers8]. The array was interfaced, via a thin hydrogel layer (0.5 mm

hickness), with the underlying tissue (skin, fat and muscle thick-esses of 2, 3 and 100 mm and resistivities of 909, 43 and 3.3 � mespectively). Exemplar tissue thickness values were taken from anRI scan of an individual’s leg although these thicknesses will vary

ig. 3. Finite element model constructed in SEMCAD, consisting of an 8 × 8 metallectrode array with remote anode interfaced with the skin by a 0.5 mm thick hydro-el layer.

model (7.1 mm nerve depth, hydrogel resistivity 1000 � m). The absolute values forthis normalisation point (which can be used to derive values for all the other pointin V/mm per mA) were 1.414 V/mm for a total drive current of 1.62 mA.

between different people. Tissue resistivity has been estimatedfrom Gabriel et al. [13] combined with values used in a similar appli-cation of FES FE modelling in reference [11]. The anode electrodewas 100 mm from the edge of the electrode array to emulate typicalanode–cathode spacing in drop-foot applications. Fig. 3 shows theFE model constructed in SEMCADv14.2.

The model was used to look at the effect of changing the resis-

Fig. 5. Focality of the electrode array measured from the area at a specified depthgreater than 50% of the maximum electric field gradient of the model.

970 G. Cooper et al. / Medical Engineering & Physics 33 (2011) 967– 972

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ig. 6. Example plots at various depths for a VE consisting of 4 × 4 electrodes for hilectrodes but at different depths. (a) Low resistivity hydrogel interface at a tissue desistivity hydrogel interface at a tissue depth of 10 mm. (d) High resistivity hydroepth of 25.6 mm. (f) High resistivity hydrogel interface at a tissue depth of 25.6 mm

The model was used to calculate the activation function for atraight nerve at three depths, defined as the electric field gradi-nt along the x direction of the model (i.e. parallel to the nervehich is assumed to run from anode to cathode) [14], at depths

f 7.1, 10.1 and 25.6 mm below the skin surface chosen to repre-ent a nerve running parallel with the surface of the model, parallelo the anode–cathode axis, and at superficial, medium and deepepths. ‘Virtual electrodes’ (a group of driven array electrodes) ofize 1, 4, 9 and 16 electrodes (1 × 1, 2 × 2, 3 × 3 and 4 × 4 respec-

right) and low (on left) resistivity hydrogel. The data is plotted in the plane of thef 7.1 mm. (b) High resistivity hydrogel interface at a tissue depth of 7.1 mm. (c) Lowerface at a tissue depth of 10 mm. (e) Low resistivity hydrogel interface at a tissue

tively) for both high and low hydrogel resistivity were driven withan approximation to a constant current source applied to each arrayelectrode.

3.2. Stimulation efficiency and focality results

Stimulation efficiency was defined as the calculated maximumelectric field gradient divided by the stimulation current applied,and normalised to that of the maximum efficiency model (1 × 1

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ctivated electrodes, 7.1 mm depth and 1000 � m hydrogel inter-ace layer). The normalised efficiency results are plotted in Fig. 4,he graph shows efficiency against the number of activated elec-rodes for different nerve depths and hydrogel interfaces (high andow resistivity).

The focality (focus area) of the stimulation was also calculatedrom the electric field gradient (defined here as the area of the elec-ric field gradient at a specified depth that was above an arbitrary0% of the maximum electric field gradient at that depth).

The results for the focus area for all the models and depths arehown in Fig. 5.

As an example of the focality plots so obtained, Fig. 6 shows thectivation function for a virtual electrode (VE) consisting of 16 acti-ated electrodes (4 × 4) with both high and low resistivity hydrogelt three different depths.

. Discussion

Fig. 2 shows that all four hydrogel samples exhibited a largerop in resistivity over the course of 7 days to levels comparableith conventional, low resistivity hydrogel. It is believed that this

hange was a result of ions in sweat from the skin surface passingnto the hydrogel layer thus decreasing its electrical resistivity.

Stimulation efficiency and focus area are both affected by hydro-el resistivity. The efficiency plot, Fig. 4 shows that, for a fixedrive current to the virtual electrode (current shared equally by alllectrodes), a virtual electrode comprising just a single electroderovides the greatest efficiency under all the tested conditions.his is because the current density underneath it is high, result-ng in a relatively high local electric field gradient. The efficiencys decreased for the low resistivity hydrogels relative to their highesistivity counterparts is partly due to the hydrogel ‘wicking’ cur-ent to undriven electrodes surrounding the virtual electrode (theseetal structures then spread the current still further) and partly to

he spread of current in the hydrogel that would occur even in thebsence of undriven electrodes.

The focus area curves, Fig. 5, are not simple to interpret becausehere are additional mechanisms occurring. They show two generalnd expected trends, namely the increase of focus area with theumber of active electrodes and with decreasing hydrogel resistiv-

ty (the increase in focus area with decreasing hydrogel resistivitys also shown in Fig. 6). For the virtual electrode consisting of aingle electrode, the activating areas for each of the six conditionsested (3 depths and two hydrogel types) all produce single poolshich occur under the electrode edge nearest the anode. But for 4,

and 16 active electrodes at the most superficial depth (7.1 mm)nd with the high resistivity hydrogel show 2, 3 and 4 discreteools respectively of activation under the electrodes along the x-xis. An example of this is shown in Fig. 6b for 16 active electrodes.nder all other conditions the activation area is a single pool cen-

ered approximately on the virtual electrode. In general the effectf hydrogel resistivity is less in these circumstances although thexpected general trend of area increasing with decreased resistivityan be observed in most cases (Fig. 6).

Current spread within the tissue increases and hence focality oftimulation decreases significantly for the lower resistivity hydro-el. This is caused by transverse current flow in the hydrogel andurrent flow across neighbouring passive electrodes.

The upper limb model proposed by Kuhn et al. [11] explored theffects of hydrogel resistivity on stimulation efficiency. Their mod-ls also included different electrode array geometries and high and

ow skin impedances. Unfortunately it is not possible to directlyompare their results with the ones in this paper as their electrodefficiencies are reported as increases in required stimulation cur-ent but they have not reported their initial stimulation currentbove which these increases occur.

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& Physics 33 (2011) 967– 972 971

A limitation of this research is that the hydrogel resistivity datawas only measured on a single test subject. However, it is expectedthat similar results would be observed if multiple subjects had beenused (regardless of the demographic). This research highlights theproblems with using a hydrogel skin interface with electrode arraysand further work is required to investigate practical solutions tothis problem.

We conclude that using a continuous layer of high resistivityhydrogel as an interface between an electrode array and the skin isa good solution for short periods of time (perhaps less than a cou-ple of hours). Efficiency of stimulation reduces as the resistivity ofthe hydrogel layer is decreases and the stimulus current becomesmore diffuse. This will result in the need to apply higher electri-cal stimulation currents to create a functional response, which willpotentially increase discomfort during stimulation. These highercurrents would also require greater electrical energy to be sup-plied which will, in turn, reduce battery life. Focality of stimulationis also reduced with a reduction in resistivity of the hydrogel layeralthough this effect is reduced with depth in the tissue. Hence elec-trical stimulation of deep nerves (∼25 mm or deeper) will havesimilar stimulation focus areas when using both high and low resis-tivity hydrogel layers.

Our modelling results suggest that there are benefits of usinga high resistivity hydrogel interface layer over a low resistivityhydrogel interface layer but our experimental results suggest thatthis benefit would be lost following more than ∼1 day of continuouswearing due to a reduction in resistivity of the hydrogel.

Acknowledgements

The work was funded by the UK NIHR (HTD480). The viewsexpressed in this publication are those of the author(s) and not nec-essarily those of the NHS, the National Institute for Health Researchor the Department of Health.

Conflict of interest statementNone of the authors have any conflict of interest related to this

work.

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