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Keywords:

  • organic bioelectronics;
  • conducting polymers;
  • cutaneous recordings;
  • healthcare materials;
  • ionic liquids

Cutaneous electrophysiological recordings are used in standard clinical tests to measure the integrity of an organ or a specific physiological function. For example, electroencephalography (EEG) is a diagnostic tool, which typically uses metallic electrodes on the surface of the scalp to measure electrical activity in the brain.[1] When compared with other neurolo­gical diagnostic methods, EEG is less invasive and more cost-effective. For this reason, EEG continues to be the method of choice for clinicians when testing for neural pathologies.[2] It is often combined with other electrodes placed on the skin to simultaneously or independently measure cardiac activity (electrocardiography)[3] or muscular response (electromyography[4]–electrooculography).[5] Despite their widespread use in the majority of electrophysiological diagnostic procedures, currently used cutaneous electrodes have some shortcomings, especially in long-term measurements:[6] Temporal stability is an important characteristic for diagnostic tools, and in many cases, cutaneous recordings need to be performed over several days.[7] Ag/AgCl electrodes, which are the current gold standard in cutaneous recordings, require a liquid electrolyte to decrease the electrode/skin impedance. This is undesirable for several reasons: First, the electrolyte often dries out over the course of only a few hours when exposed to open air.[8] As a result, the impedance of these electrodes usually increases and their ability to record meaningful signals is lost. Second, in cases where high-density recordings are necessary, as in some EEG helmets, short circuits can occur if the liquid electrolyte leaks between two adjacent electrodes.[9] Third, adding the electrolyte gel to each electrode is time-consuming and causes discomfort to both the patient and the caregiver. For these reasons, there is significant motivation to develop alternative electrolytes for use in cutaneous measurements.

Much effort has been put into improving dry electrodes, which do not utilize a liquid electrolyte. One study reports the use of only a very small volume of gel, which is released by pressure on the electrode, but this requires a system, which is able to apply pressure to all the electrodes.[10] Several studies report the fabrication of microstructures on the surface of the electrodes in order to increase the surface area and decrease the electrode impedance.[11-14] We recently reported the use of conducting polymers as dry electrodes with enhanced performance.[15] Another solution involves the use of spring-loaded contacts embedded in a flexible substrate.[8] Although these are promising ideas, the lack of an electrolyte means that movement artifacts are more likely to occur during recording.

Ionic liquids (ILs) are salts that are in a liquid state at room temperature.[16] They have recently inspired a great deal of research spanning a number of different fields due to their high chemical and thermal stability as well as their excellent ionic conductivity. ILs have been used in many biological applications, including as solvents that enhance the stability of proteins.[17] Although several IL families showing low cytotoxicity have been identified,[18-20] the establishment of general design rules for IL biocompatibility is the focus of ongoing work. ILs can be polymerized to yield gels, which are especially appealing for use as a quasi-solid-state electrolytes.[21] Their lack of leakage makes them an ideal candidate for many electrochemical applications,[22] including bio-sensing electrodes,[23] and transistors.[24] IL gels seem to be an appealing option for use with cutaneous electrodes because they do not leak or dry out and can be integrated with devices during fabrication, thereby addressing all short comings of electrodes currently used in electrophysiological recordings.

In this communication, we report the use of IL gel-assisted electrodes in long-term cutaneous recordings. We incorporated the IL gel onto electrodes made of Au and of the conducting polymer poly(3,4-ethylenedioxythiophene) doped with poly(styrene sulfonate) (PEDOT:PSS). The latter was used as it was shown to yield high-performance dry electrodes.[15] The electrodes were deposited on a thin film of parylene C to render them conformable to the skin. We show that the IL gel improves the performance of the electrodes and helps maintain a low impedance over longer periods of time than commercially available electrodes.

Figure 1 illustrates the fabrication process for the devices. Gold electrodes were patterned using shadow mask evaporation onto a 2-μm thick layer of parylene C. We then insulated the electrodes with two additional layers of parylene, with a thickness of 2 μm each. An anti-adhesive soap was included between these two layers, to allow the top layer to be peeled-off at a later stage. Subsequently, a standard photolithographic process was used to expose only the electrode sites to etching. Once the devices were etched down to the gold layer, the conducting polymer PEDOT:PSS was spun over the entire wafer. The top layer of parylene was peeled-off, patterning the PEDOT:PSS film over the electrode area and leaving the parylene layer insulating only the interconnects (the details of this patterning technique were reported in the literature).[25] This process yields electrodes with a total thickness of only ≈4 μm, a size which endows them with a high conformability. A support foil was laser-cut out of Kapton and gold was deposited on it. It was connected to the electrode using a small amount of adhesive placed on its periphery. Figure 1a shows how the electrode and support foil were attached to each other. Figure 1b shows how this fabrication process resulted in a conformal 4 μm thick electrode, which was also strong enough to withstand connection. The gel was made using the IL (1-ethyl-3-methylimidazolium ethyl sulfate), the polymer poly(ethylene glycol) diacrylate, and a photoinitiator.[24] It was polymerized directly on the electrode during the fabrication process. A variety of electrodes were prepared using this general process (Figure 1c), including dry Au and PEDOT:PSS electrodes, Au and PEDOT:PSS electrodes with the IL gel, and a commercial Ag/AgCl electrode (which included an aqueous gel).

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Figure 1. A) Schematic of the electrode assembly, and B) cross-section of an electrode. C) The different electrode configurations tested. D) Schematic of the electrode positions on a subject's arm. The working electrode (W.E.) and counter electrode (C.E.) were placed on the forearm, 5 cm away from each other. The reference electrode (R.E.) was placed on the arm, 30 cm away from the W.E.

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In order to compare the performance of these electrodes in cutaneous recordings, we tested the electrode/skin impedance on a healthy volunteer. The electrodes were placed on the subject's arm as shown in Figure 1d using a three-electrode configuration, where both counter and reference electrodes were commercial Ag/AgCl electrodes. Figure 2 shows typical electrical impedance spectra for all electrodes tested. In terms of dry electrodes, PEDOT:PSS demonstrated a lower impedance than Au, consistent with results obtained in microelectrode arrays in vitro.[25] Adding an IL gel decreases the impedance of both electrodes considerably: At 1 kHz, which is the frequency at which EEG electrodes are tested clinically, the impedance of the IL gel-assisted and aqueous gel-assisted electrodes is similar (see below for measurement error).

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Figure 2. Electrical impedance spectra corresponding to dry Au and PEDOT:PSS electrodes, IL gel-assisted Au and PEDOT:PSS electrodes, and a commercial Ag/AgCl electrode with an aqueous gel, all in contact with skin.

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The temporal stability of the gel-assisted electrodes is shown in Figure 3. The same setup as for the previous measurement was used, and impedance was measured at 1 kHz on a healthy volunteer's arm over a period of 3 d. All three electrodes began in approximately the same range, but while the impedance of the commercial electrode steadily increased, the impedance of both IL gel-assisted electrodes remained relatively constant. Both IL gel-assisted electrodes were able to record with a low impedance over 3 d, while the impedance of the commercial electrode increased dramatically after only 1 d. In fact, after 20 h, the Ag/AgCl electrode showed an impedance that was too high for high-quality cutaneous recordings and recordings from this electrode were stopped.

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Figure 3. Electrode/skin impedance measured at 1 kHz for IL gel-assisted Au and PEDOT:PSS electrodes and commercial Ag/AgCl electrode with an aqueous gel.

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Although both IL gel-assisted electrodes showed similar values of impedance, the PEDOT:PSS electrodes performance exhibited less variation. The error bars in Figure 3 represent measurements recorded by five different devices for each type of IL gel-assisted electrode. The observed variations are due to a multitude of reasons, including device-to-device variations ­(fabrication, electrode work-function), and changes in the hydration of the skin. The conducting polymer-based electrode is shown to exhibit less variation than the Au-based one, owing to a more stable surface. This is consistent with work on implantable microelectrodes, which showed that conducting polymer electrodes offer lower drift in vivo than electrodes made from metals.[26, 27] As the measurements were carried out on a person's arm, many factors caused slight changes in the performance of the electrodes. After the subject showered, for example, the electrodes were partially rehydrated and the impedance decreased. As it can be seen in Figure 3, both IL gel-assisted electrodes followed the same cycle: Lowering of impedance due to rehydration and subsequent increasing of impedance during the course of the day. This variation is due to a measurement protocol that represents an accurate model for cutaneous recordings under realistic conditions, as the latter are carried out on patients who inevitably sweat, particularly if bandages are used.

In this work, we used a photo-crosslinkable IL gel to show that these materials help remedy the most important shortcoming of cutaneous electrodes, namely they can be used for long-term measurements. Moreover, as IL gels do not flow, they can enable high-density electrode arrays without causing short circuits. Finally, IL gels can be incorporated onto the electrodes during fabrication, yielding ready-to-use devices. The fabrication scheme outlined here is compatible with such an endeavor, and the IL gel can be patterned at the same time as the conducting polymer using the peel-off step. There exists a great variety in IL structures and in ways to render them in a gel form. Future research should focus on exploring the tradeoff between the conductivity and mechanical properties of these materials to deliver IL gel-assisted electrodes with even higher performance.

In conclusion, we demonstrated that ionic liquid gel-assisted electrodes can be used for long-term cutaneous recordings. We fabricated conformal electrodes made of Au and PEDOT:PSS and compared their performance in a dry state and in conjunction with an IL gel. The IL gel decreases impedance at the interface with human skin to levels that are similar than those of commercial electrodes (at 1 kHz). The IL gel did not dry out and the electrodes continued to show a low impedance over the course of 3 d. The commercial electrode, on the other hand, gave up after only 20 h. The IL gel-assisted electrodes provide a means of recording cutaneous electrophysiological data over extended periods of time, which is often a necessity during diagnostic procedures. As such, they provide a path towards accurate and stable electrophysiological recordings, meaning clinicians could rely less heavily on more invasive diagnostic techniques.

Experimental Section

  1. Top of page
  2. Experimental Section
  3. Acknowledgements

Device Fabrication: Parylene C was deposited on a silicon wafer in a SCS Labcoater 2 to a thickness of 2 μm (at which thickness parylene films are pinhole free). Subsequently, 10 nm of chromium and 100 nm of gold were deposited using a metal evaporator. The metal was evaporated only onto the shape of the 8 mm diameter electrode using a shadow mask. This was followed by the deposition of an additional 2 μm thick layer of parylene, a non-adhesive layer, and a last 2 μm thick layer of parylene. AZ9260 photoresist was spun over the wafer and exposed to UV light using a SUSS MBJ4 contact aligner and a shadow mask. Following development of the photoresist using AZ developer, two layers of PEDOT:PSS were spun onto the wafer. The first layer was spun at 1500 rpm for 30 s; the second at 600 rpm for 30 s. The PEDOT:PSS was prepared by mixing 45 mL of aqueous dispersion (PH-1000 from H.C. Clark) with 5 mL of ethylene glycol, 6 drops of dodecyl benzene sulfonic acid (DBSA), and 0.5 mL of 3-glycidoxypropyltrimethoxysilane (GOPS, as a cross-linker). The PEDOT:PSS film on the wafer was baked at 140 C for 30 min and subsequently immersed in deionized water to remove any excess low-molecular-weight compounds. The top layer of parylene was then removed, leaving behind the PEDOT:PSS only on the electrodes and contact pads. The remaining parylene C served to insulate the interconnects. A Au-coated kapton support foil was glued to the device to give it some rigidity and enable an easy and stable connection to the recording system.

IL Gel Preparation: The IL gel was prepared by mixing the IL (1-ethyl-3-methylimidazolium ethyl sulfate) and the polymer poly(ethylene glycol) diacrylate (30% of the total weight solution). We then added 0.3% of the photoinitiator 2-hydroxy-2-methylpropiophenone. We mixed until no phase separation was visible. Ten microliters of the solution was pipetted onto the electrode, and the gel was cross-linked using exposure to UV light for approximately 1 min.

Impedance Measurements: Informed signed consent was obtained from the healthy volunteer on whom the impedance was measured. Impedance spectra were measured using an Autolab potentiostat equipped with an FRA module, with the electrodes placed on the subject's arm. The applied voltage was 0.01 V. Dry and IL gel-assisted electrodes were taped to the skin using surgical tape. Commercially available Ag/AgCl electrodes (Comepa Industries) for ECG were used (as received) as the aqueous gel-assisted working electrodes, and as the counter and reference electrodes for the impedance measurements. The latter were located 5 cm and 30 cm away from the working electrode, respectively. For the stability measurements shown in Figure 3, fresh counter and reference electrodes were used for each measurement. For all measurements and for all electrode types, we waited 10 min after the application of a fresh electrode before a measurement was taken.

Acknowledgements

  1. Top of page
  2. Experimental Section
  3. Acknowledgements

P.L. and C.J. contributed equally to this work. This work was supported though grants by the ANR, region PACA, and MicroVitae Technologies. The prototyping and fabrication of the electrodes were performed at the Centre Microelectronique de Provence. J.R. acknowledges a Marie CURIE postdoctoral fellowship. C.J. was an iREU student of the National Nanotechnology Infrastructure Network. The authors would like to thank Vincenzo F. Curto from DCU for his help on the ionic liquid gel formulation, and Jean-Michel Badier and Christian Benar from INSERM for fruitful discussions.