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A PDMS-Based Integrated Stretchable Microelectrode Array (isMEA) for Neural and Muscular Surface Interfacing Liang Guo, Member, IEEE, Gareth S. Guvanasen, Student Member, IEEE, Xi Liu, Christopher Tuthill, T. Richard Nichols, and Stephen P. DeWeerth, Senior Member, IEEE
Abstract—Numerous applications in neuroscience research and neural prosthetics, such as electrocorticogram (ECoG) recording and retinal prosthesis, involve electrical interactions with soft excitable tissues using a surface recording and/or stimulation approach. These applications require an interface that is capable of setting up high-throughput communications between the electrical circuit and the excitable tissue and that can dynamically conform to the shape of the soft tissue. Being a compliant material with mechanical impedance close to that of soft tissues, polydimethylsiloxane (PDMS) offers excellent potential as a substrate material for such neural interfaces. This paper describes an integrated technology for fabrication of PDMS-based stretchable microelectrode arrays (MEAs). Specifically, as an integral part of the fabrication process, a stretchable MEA is directly fabricated with a rigid substrate, such as a thin printed circuit board (PCB), through an innovative bonding technology—via-bonding—for integrated packaging. This integrated strategy overcomes the conventional challenge of high-density packaging for this type of stretchable electronics. Combined with a high-density interconnect technology developed previously, this stretchable MEA technology facilitates a high-resolution, high-density integrated system solution for neural and muscular surface interfacing. In this paper, this PDMS-based integrated stretchable MEA (isMEA) technology is demonstrated by an example design that packages a stretchable MEA with a small PCB. The resulting isMEA is assessed for its biocompatibility, surface conformability, electrode impedance spectrum, and capability to record muscle fiber activity when applied epimysially. Index Terms—Compliant, epimysial, integrated packaging, microelectrode array (MEA), microfabrication, neural prosthesis, neural recording and stimulation, polydimethylsiloxane (PDMS), stretchable, surface. Manuscript received December 15, 2011; revised February 16, 2012; accepted March 16, 2012. Date of publication May 08, 2012; date of current version March 07, 2013. The isMEA development work was supported by the U.S. NIH Grant EB006179. The cat experiment was supported by the U.S. NIH Grant HD32571. This paper was recommended by Associate Editor S. Leonhardt. L. Guo and C. Tuthill are with the Wallace H. Coulter Department of Biomedical Engineering, Georgia Institute of Technology and Emory University, Atlanta, GA 30332 USA (e-mail:
[email protected]). G. S. Guvanasen is with the School of Electrical and Computer Engineering, Georgia Institute of Technology, Atlanta, GA 30332 USA. X. Liu is with the George W. Woodruff School of Mechanical Engineering, Georgia Institute of Technology, Atlanta, GA 30332 USA. T. R. Nichols is with the School of Applied Physiology, Georgia Institute of Technology, Atlanta, GA 30332 USA. S. P. DeWeerth is with the Wallace H. Coulter Department of Biomedical Engineering, Georgia Institute of Technology and Emory University, Atlanta, GA 30332 USA, and also with the Department of Biomedical Engineering, Khalifa University of Science, Technology, and Research, Abu Dhabi, United Arab Emirates (e-mail:
[email protected]). Color versions of one or more of the figures in this paper are available online at http://ieeexplore.ieee.org. Digital Object Identifier 10.1109/TBCAS.2012.2192932
I. INTRODUCTION
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EURAL recording and stimulation hold promise for both advancing our understanding of the underlying physiological mechanisms and providing unique diagnostic and therapeutic solutions to neurological disorders and disabilities. Communications between neural tissues and electrical hardware are made possible through electrodes that transduce ion-based currents in the neural tissue to electron-based currents in the electrical circuit during recording, and vice versa during stimulation. Historically, single rigid-needle electrodes were the primary tool for interfacing with the neural systems [1]. To increase signal dimensionality, multiple-electrode approaches were first developed through handcrafted microwire bundles. However, it was not until semiconductor microfabrication technologies were applied to the fabrication of microelectrode arrays (MEAs) that the field of neural interfacing started to thrive owing to the significant improvements on device resolution and reliability. These rigid-needle MEAs are typically fabricated using silicon, glass, and metals [2]–[8]. A significant problem with the rigid-MEA approaches has been their mechanical impedance mismatch with the surrounding tissues. Besides causing trauma during insertion, these devices often fail in the long run due to intensive scar-tissue encapsulation induced by micro-motions between the hard materials and surrounding tissues [9]. Thus, soft MEAs are developed more recently using polymeric substrates such as polyimide [10]–[16], parylene [17], SU-8 [18], etc. However, [10], [17] and with Young’s moduli of [19], respectively, these materials are still about four orders of magnitude stiffer than nerves (Young’s modulus [20]) and six orders of magnitude stiffer than brain [1]) (see Fig. 1). tissue (Young’s modulus To further reduce the invasiveness of penetrating MEAs, an alternative approach can be employed by using MEAs that interface only with the tissue surface. This approach has worked effectively in applications where reduced accessibility is an acceptable trade-off and where the functional structures of interest are located at or close to the tissue surface, such as electrocorticogram (ECoG)-based brain-computer interfaces (BCIs) [21]–[23] and retinal prostheses [24]–[28]. Indeed, some of the aforementioned polymer-based flexible MEAs are designed as thin-film devices for surface applications [15]–[17]. Although these polymeric substrates exhibit a certain level of flexibility, they are not an ideal solution to the challenge of chronic surface recording and stimulation, primarily because these flexible
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IEEE TRANSACTIONS ON BIOMEDICAL CIRCUITS AND SYSTEMS, VOL. 7, NO. 1, FEBRUARY 2013
Fig. 1. Young’s modulus spectrum of common MEA substrate materials and soft tissues.
MEAs cannot conform to the complex tissue surfaces (unless using an ultra-thin film combined with a special mesh structure [16], which, however, would reduce the device’s durability and limit the interconnect density). Being a truly compliant material with a Young’s modulus of [29], polydimethylsiloxane (PDMS, i.e., silicone rubber) offers excellent potential as a substrate material for such neural interfaces [29]–[32]. When used as a surface interface, the PDMS-based device has the capability of conforming much better to the complex tissue surfaces to form a uniform and tight contact. Application of compliant MEAs to neural surface interfacing could provide a novel and potentially powerful method for both basic neuroscience research and neural prosthetics. However, fabrication of electrical functionalities on PDMS substrates has proven very challenging, particularly for achieving high-density interconnects and integrated packaging, as are frequently required in neuroscience research and neural prostheses (e.g., a high-resolution retinal prosthesis requires at least 600 ~ 1000 microelectrodes in an area [24]). We have previously developed a high-density interconnect technology on PDMS substrate [33], [34]. In this paper, we report a PDMS-based integrated stretchable MEA (isMEA) technology whose fabrication process is compatible with the high-density interconnect technology. Thus, combining these two technologies will provide a high-resolution, high-density integrated system solution for neural and muscular surface interfacing. In this paper, we demonstrated this isMEA technology through an example design that packaged a stretchable MEA with a small printed circuit board (PCB). We then performed a set of experiments to characterize the resulting isMEA. We used rat cortical neuronal culturing to evaluate its cytotoxicity, a curved surface to reveal its conformability, finite element modeling to assess its stretchability, impedance spectroscopy to measure its electrode impedance, and epimysial electromyography (EMG) to test its recording capability. Some preliminary results of this study have been reported previously as conference proceedings [35]–[37]. II. METHODS We begin with an explanation of the isMEA fabrication method, then we describe the experimental methods used for characterizing the resulting isMEA. A. Fabrication Method PDMS-based stretchable MEAs are usually fabricated as a sandwiched structure [29]–[32], where thin-film interconnects
Fig. 2. Cross-sectional illustration of the isMEA fabrication processes (not to scale).
made of a noble metal, such as platinum or gold, are patterned as electrodes and leads and embedded between two PDMS layers: a base layer and an insulation layer with holes to expose electrodes and/or contact pads. Our PDMS-based isMEA technology adopts the same sandwiched structure and can be easily customized to designs for various neural surface interfacing applications. Below, we describe the fabrication of an isMEA, which was designed to be packaged through a small PCB for epimysially interfacing with a feline muscle in an acute experimental setup. Please refer to Fig. 2 for an illustration of the fabrication processes. 1) Materials: The small PCB used in the isMEA was made of FR4 (250 thick) with emersion-gold traces ( thick) plated on both sides. Alignment markers on the PCB were designed as isolated metal features. The glass fabrication carrier was a 75 50 1 mm microscope slide (Fisher Scientific). PDMS (Sylgard 184, Dow Corning) was used for both the device substrate and insulation layers. The elastomer base was mixed with the curing agent at 10 : 1 weight ratio, and the mixture was left at room temperature for a minimum of 40 minutes to allow air bubbles to escape. 2) Preparation: The via-bonding process [34] requires the bonding substrate (i.e., the small PCB) to be in the same plane as the fabrication substrate (i.e., the PDMS base layer). So, the first step of fabrication is to prepare the sample to form a coplanar structure [Fig. 2(a)]. First, an anti-adhesion layer (10 Å Ti/50 Å Au, both deposited at 3 Å/s in an e-beam evaporator) was coated on the glass slide to facilitate release of the final device from the glass carrier. PDMS prepolymer was spin-coated on the glass slide at 450 rpm for 30 s to yield a thickness matching that of the PCB. A cleaned PCB was gently dropped into the uncured prepolymer, and the sample was left on a level surface at room temperature for at least 12 hours before fully curing the PDMS on a 60 hotplate for 1 hour followed by 80 oven baking for another hour. The 12 hours at room temperature allowed the PCB to settle down in
GUO et al.: A PDMS-BASED isMEA FOR NEURAL AND MUSCULAR SURFACE INTERFACING
the prepolymer without trapping air bubbles around its edges, a by-product of fast curing. This cured sample formed a coplanar fabrication substrate [Fig. 2(a)] for the subsequent via-bonding process. It was found necessary to cover the backside of the PCB with a piece of Kapton tape (25 thick) to prevent the PCB vias from drawing prepolymer up to the top surface by capillary force. It was also found that pressing on the top surface of the PCB squeezed out the air bubbles underneath, which helped the PCB to settle properly in the prepolymer. 3) Via-Bonding and Interconnect Patterning: The via-bonding process included Fig. 2(b) – (e), and the gold interconnect patterning using an SU-8 lift-off method [33] was combined in the last step [Fig. 2(e)] to define the electrodes and leads at the same time. Tapered sacrificial posts made of negative photoresist NR5-8000 (Futurrex, Inc.) were patterned on each contact pad of the PCB where a via-bond was to be formed, using a projection UV lithography method [34] [Fig. 2(b)]. The height of the sacrificial posts defined the coating thickness of the following PDMS layer. 400 rpm/30 s was used to result in a sacrificial post height of [38]. If a thinner device is desired for better conformability, an anti-adhesion layer (10 Å Ti/50 Å Au), which facilitates easy separation of the two PDMS layers at the interface, can be partially coated onto the PDMS surface (i.e., except on areas surrounding the PCB) before applying the second PDMS layer, as shown in Fig. 2(c). Without using this anti-adhesion layer, the following PDMS layer will be polymerized with the substrate PDMS as a whole. In Fig. 2(c), immediately following a brief oxygen plasma treatment to activate the sample surface, a PDMS layer was spin-coated. The spin-coating recipe for the PDMS layer depended on height of the sacrificial posts, and the thickness of the PDMS needed to be made slightly lower than the height of the sacrificial posts. The sample was then left on a level surface for 1 hour, baked on a 60 hotplate for 1 hour, and finally placed in an 80 oven for at least 4 hours until the new PDMS layer became fully cured. To remove the sacrificial posts, the sample was first briefly etched in a Plasma Thermal Reactive Ion Etcher (RIE) system to remove any potential PDMS residues covering the sacrificial posts [34]. Following this step, the sacrificial posts were dissolved by immersing the PCB in acetone for a few minutes. Any NR5-8000 residues in the inclined-vias were removed using an RIE de-scum process ( / ). Contact pads on the PCB reserved for final assembly were also opened during the same processes [Fig. 2(d)]. An SU-8 lift-off method [33] was then used to pattern the interconnects and to complete the via-bonding process [Fig. 2(e)]. The gold film deposited on the slope of the inclined-vias bridged the interconnects on the PDMS with those on the PCB [see Fig. 4(b)]. By iterating this via-bonding process [Fig. 2(b) – (e)], multilayer interconnected interconnects can be achieved subsequently to facilitate wiring of higher density designs [34]. 4) Passivation: The processes for passivation of the device [Fig. 2(f) and (g)] were essentially the same as those for making the inclined-vias [Fig. 2(b) – (d)]. Processing parameters were tuned to achieve a desired insulation thickness (e.g., 10 ). In
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addition, by manipulating the associated processing parameters, different electrode profiles can be fabricated, including conicalwell electrodes [38], [39]. 5) Post-Fabrication Assembly: After completion of the fabrication, the contour of the isMEA was cut out first using a razor blade, and the isMEA along with the thick PDMS substrate was peeled off the glass carrier. The thin device body was separated from the thick PDMS substrate by peeling using a pair of tweezers, and the excess PDMS substrate was cut off [Fig. 2(h)]. Then, wires were soldered on the PCB, and finally the PCB area was encapsulated on both sides by additional PDMS. B. Dissociated Neuronal Culturing Rat/mouse cortical neuronal culturing is a standard assay for evaluating cytotoxicity of a drug or biomaterial. Here we adopted this assay to test the cytotoxicity of our isMEA as fabricated. Cortical neurons were derived from embryonic day 17–18 rat fetuses by isolating the cerebral cortices, and dissociating them with trypsin (0.25%) + 1 mM EDTA (10 minutes at 37 ) followed by DNase (0.15 mg/mL). These neurons were plated in culture dishes at densities of approximately 1 and 5 , to form high and low concentration cultures respectively. Before neuronal plating, an isMEA sample had been treated with polyethyleneimine (PEI), thoroughly washed, fixed to the bottom of each culture dish, and coated with laminin. The cultures were fed neuronal medium (Neurobasal medium + 2% B-27 + 500 L-glutamine) and maintained in a tissue culture incubator (35 , 5% , 9% ) for 2–3 weeks. C. Finite Element Modeling of the isMEA as an Epimysial Interface To simulate the epimysial mechanical performance of the isMEA, we constructed a three-dimensional finite element model using Abaqus FEA (ABAQUS Inc.) running on a 32 GB memory server. As shown in Fig. 3, the isMEA head was laid with electrodes facing down on a PDMS stretching substrate (210 124 10 mm) to mimic an epimysial interface. All of the dimensions of the isMEA head were identical to those of a real device. Specifically, the isMEA head model consisted of a 50 thick PDMS base layer, 300 nm thick gold interconnects, and a 10 thick PDMS insulation layer. The electrode openings were 1 mm in diameter. The PDMS stretching substrate was designed to serve as a simple model of a muscle surface, to which a 5% uniaxial strain, either longitudinal or transversal, was applied to simulate muscle stretching. A viscoelastic model [41] was used for the PDMS material; and an elastoplastic model [42] was used for the thin-film gold. The ultra-thin Ti adhesion layer (30 nm thick) between the gold film and the PDMS substrate in the actual isMEA was neglected to reduce difficulty and computational complexity of the simulation. It was assumed that all materials were isotropic and that the interfacial bonding between the isMEA and the PDMS stretching substrate was perfect. Two different sets of boundary conditions, longitudinal ( direction) and transversal ( direction), were applied to this
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Fig. 3. Finite element model for simulation of the isMEA’s mechanical performance as an epimysial interface. The isMEA head was laid with electrodes facing down on a 1 cm thick PDMS stretching substrate to mimic the epimysial interfacing application. Five percent uniaxial strain was applied to the PDMS substrate either longitudinally or transversally.
model to study the isMEA’s mechanical performance under uniaxial loading conditions. When applying a 5% strain, either longitudinally or transversally, the PDMS stretching substrate was fixed on one side, and a displacement was uniformly applied on the surface of the opposite side. The resulting strain in the gold interconnects was then analyzed. D. Electrode Impedance Spectroscopy Electrode impedance spectra of the isMEA were measured using a spectrum analyzer (SRS Dynamic Signal Analyzer, SR785) with a two-terminal cell setup. A custom-made comparing circuit was used to interface the isMEA to the spectrum analyzer. The isMEA was laid on a flat surface. One droplet of saline solution (HBSS 1X, Gibco) was applied to the electrode surface. A silver wire connected to one probe was dipped into the solution droplet, without touching the electrode surface. The other probe was hooked on the corresponding connector pin to close the circuit. A frequency sweep was applied from 5 Hz to 20 kHz with a voltage amplitude of 50 mV. The magnitude of impedance spectra of 16 recording electrodes and one reference electrode were recorded. E. Multichannel Epimysial Recording All experimental procedures within this study were conducted in accordance with the guidelines of the National Institutes of Health and the Georgia Tech Institutional Animal Care and Use Committee. In the experiment, the cat was tracheostomized to control for its isoflurane anesthetic levels. Under deep surgical anesthesia, the animal was subject to a precollicular decerebration, and all brain tissue rostral to the transection was removed. The tendon of the right limb medial gastrocnemius (MG) muscle was detached from the insertion and clamped while the right knee was fixed. This muscle preparation was performed to reduce the introduction of movement artifacts during the recording of EMG activity. A bipolar electrode was implanted adjacent to the left tibial nerve for the purpose of electrically inducing a crossed extensor reflex, and
Fig. 4. A fully assembled 16-channel isMEA. (a) The blue band was residue of the thick PDMS substrate on which an anti-adhesion layer (gold films thinner than 60 Å appear blue [40]) had been coated. The excess thick PDMS substrate . Inset, had been cut off. The total thickness of the isMEA body was a close-up view of an electrode, of which wrinkles were displayed in the gold film/PDMS structure. (b) A close view of the PCB area. Wires were soldered onto the PCB pins, followed by PDMS encapsulation of the PCB on both sides. The two right-angled metal patterns served as alignment markers. Inset, via. bonds on the PCB contact pads. The outer diameter of the vias was
a ground electrode was implanted in the fascia of the right hind limb. Following completion of the surgery, the cat was taken off anesthetic. The animal was euthanized at the completion of the experiment with concentrated pentobarbital. The isMEA was placed upon the surface of the right MG muscle [see Fig. 10(a)] which was activated through a crossed extensor reflex by electrically stimulating the left tibial nerve with 100 /0.8 V square pulses at 40 Hz. Twelve electrodes on the isMEA were configured for bipolar recording with an electrode at the periphery used as the common reference [denoted by “ ” in Fig. 10(a)], and the isMEA cable was connected to the input terminals of a custom-built 12-channel amplifier. Recorded signals were amplified by a gain of 1000 and then digitized at a sampling rate of 3.6 kHz/channel before being sent to the computer. No real-time filtering, neither through hardware nor through software, was applied to the raw data. Post data processing was performed using Matlab 7.0 (MathWorks, Inc.). Digital filters were designed and applied to the raw data to remove noise. Specifically, a 60 Hz comb filter was used to remove the 60 Hz electromagnetic noise; a 40 Hz comb filter was used to remove the stimulation artifacts; and a fourth-order Butterworth bandpass filter (0.15 Hz ~ 1 kHz) was used to eliminate noise of lower and higher frequencies.
GUO et al.: A PDMS-BASED isMEA FOR NEURAL AND MUSCULAR SURFACE INTERFACING
Fig. 5. A prototype isMEA directly integrated with an IC chip. A 16-channel amplifier die (RHA1016, Intan Technologies, UT) was used. The die had dimen, with 85 85 square bonding sions of 3.1 mm 4.3 mm 380 pads distributed at the peripheries. Via-bonds were formed on these bonding with a minimum spacing of pads. The gold interconnects had a width of 20 . 25
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Fig. 6. Dissociated neuronal culture on isMEA surface. The isMEA samples used for testing had the same design as the one shown in Fig. 5 to take advantage of the smaller electrode area. The isMEA sample had gone through the entire fabrication processes, but without actually embedding the die in the device. (a) High-density culture. The image shows the electrode area. (b) Low-density culture. The image shows the space nearby electrodes.
III. RESULTS We demonstrated the PDMS-based isMEA technology by an example isMEA that was designed to be packaged through a small PCB for epimysially interfacing with a feline muscle in an acute experimental setup. A series of experiments was performed to characterize the resulting isMEA, including surface conformability test, mechanical performance analysis, electrode impedance spectroscopy, and multichannel epimysial recording. A. Fabrication Results The fabrication processes are illustrated in Fig. 2. A fully assembled device is shown in Fig. 4. This 16-channel isMEA had in total 17 electrodes (16 recording electrodes plus one reference electrode) of the same dimension (1 mm diameter) and one large rectangular ground electrode [right most in Fig. 4(a)]. The gold film thickness can be varied from 50 nm to 500 nm (500 nm in this specific device) with 300 nm as a trade-off between durability and flexibility. The total thickness of the isMEA body was . A close-up view of an electrode [inset, Fig. 4(a)] revealed wrinkles in the gold film/PDMS structure, which had been verified under both optical microscope and scanning electron microscope (SEM) not to be cracks. At these wrinkles, no gold film delamination was observed, rather, the underneath PDMS wrinkled together with the gold film. A close view of the PCB area [Fig. 4(b)] shows the via-bonds (inset) that bridged the gold interconnects to the PCB contact pads. On the PCB, wires were soldered onto the interface pins, followed by PDMS encapsulation of the PCB on both sides. To reveal a potential capability of our isMEA technology, we also designed and fabricated an isMEA that was directly integrated with the bare die of an integrated circuit (IC) chip (a 16-channel amplifier, RHA1016, Intan Technologies, UT). Fig. 5 shows such a prototype. Functionality of this design has not yet been tested. We will report it with its specific application to peripheral nerve interfacing in our future work. B. Dissociated Neuronal Culturing We evaluated cytotoxicity of the fabricated device by testing if any toxic chemicals from the fabrication processes were left
Fig. 7. Demonstration of surface conformability of the isMEA. The isMEA was placed on a boiled egg (shell removed). The thickness of the isMEA head . was 60
in the final device. A rat cortical neuronal culturing experiment was performed for up to three weeks. The neurons grew at a healthy rate on the isMEA surface during the testing period (Fig. 6), with extensive neurite growth [Fig. 6(b)]. This experiment proved that little to no cytotoxic chemicals were adsorbed on the device surface after fabrication. However, it was still unknown if any toxic chemical had been encapsulated in the device body, which, given a prolonged period, could potentially leak out. Future experiments are needed to test this possibility. C. Device Conformability One of the major mechanical advantages of PDMS derived from its low Young’s Modulus is that PDMS thin films (e.g., ) are highly conformable—a property highly valued in neural surface interfaces. To demonstrate superb surface conformability of the isMEA, we placed the device over a boiled egg surface. The device conformed well to the curved egg surface, as shown in Fig. 7. When this isMEA was placed on the MG muscle of a cat, the device also conformed closely to curvatures of the muscle [see Fig. 10(a)]. From a mechanics calculation [16], the minimum conformal wrapping radius for such a 60 thick PDMS-based stretchable MEA is 1.4 mm. Such conformability of PDMS, combined with its elastomeric nature, highlights the unique advantages held by PDMS-based stretchable MEAs for neural surface interfacing applications.
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Fig. 9. Electrode impedance spectra of the isMEA. The electrode diameter was 1 mm. The gold film thickness was 500 nm. Solid line, average spectrum of the 16 recording electrodes and the reference electrode; dash-dotted lines, one standard deviation from the average.
kHz (Fig. 9). This value is within the typical range of impedance magnitude for a 1 mm diameter thin-film electrode [44]. F. Multichannel Epimysial Recording
Fig. 8. Finite element modeling results of the isMEA’s mechanical performance under epimysial interfacing conditions. (a) Equivalent strains in the gold interconnects under longitudinal ( direction) loading (5% average strain on the stretching substrate). (b) Equivalent strains in the gold interconnects under transversal ( direction) loading (5% average strain on the stretching substrate). Enlarged views are provided in each figure to reveal the sites where excess strains were generated.
D. Finite Element Modeling of the isMEA as an Epimysial Interface Although PDMS films can withstand uniaxial strains greater than 200% [20], metallic interconnects usually break at strains substantially lower in magnitude [43]. To better understand the mechanical properties of our isMEA, we created a finite element model of the device to simulate its response in epimysial interfacing conditions. Fig. 8(a) and (b) show that when a 5% uniaxial strain was applied to the stretching substrate, a strain of was induced in majority portions of the gold interconnects. However, regions of large curvature or geometric transition experienced strains greater than 8%. It was also found that transversal loading (loading in the direction) induced slightly higher strains in the gold interconnects than longitudinal loading. These results are informative for guiding the interconnect layout during isMEA designing to avoid pitfalls that could diminish the device’s mechanical performance. E. Electrode Impedance Spectroscopy Electrode impedance spectra of the isMEA revealed that the electrode impedances were uniform across the frequency range of 5 Hz to 20 kHz with an average magnitude of 5.8 at 1
To evaluate the surface recording capability of the isMEA, we recorded EMG epimysially from a cat MG muscle during a crossed extensor reflex. The device conformed well to the muscle surface throughout the experiment [Fig. 10(a)]. Fig. 10(b) shows sample recordings from the first four bipolar channels, which are marked in Fig. 10(a) with the common reference electrode denoted by “ ”. A blowup of the data in Fig. 10(b) between 10 s and 10.5 s is shown in Fig. 10(c). The frequency spectrum of the first recording is shown in Fig. 10(d). The recordings had the typical characteristics of EMG signals, with the signal power accumulated primarily under 400 Hz. Preliminary analyses of both the time and frequency domains showed that the signals from each channel had different patterns, suggesting that the electric potential across the muscle surface was not uniform. Experiments are ongoing to further investigate such a spatiotemporal pattern and its relationship to the underlying physiology. IV. DISCUSSION Reliably and effectively packaging stretchable MEAs to connect with standard rigid electronics is a major challenge hindering the otherwise widespread application of such attractive soft neural interfaces. In this paper, we described a PDMSbased isMEA technology that addressed this challenge. Combined with a previously developed high-density interconnect technology, this isMEA technology facilitates a high-resolution, high-density integrated system solution for neural and muscular surface interfacing. Specific issues with the current isMEA technology are discussed below. A. Fabrication Consideration While we have moved one step forward toward integrated packaging of stretchable neural interfaces, a specific concern
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Fig. 10. Multichannel epimysial recording. (a) The isMEA was placed on the right cat MG muscle. Bipolar configurations were used with the common reference electrode marked as ” ”. (b) Sample bipolar EMG recordings from the right MG during a crossed extensor reflex (blue). The first four channels as marked in (a) are shown. The red square waves mark the stimulation periods. (c) A blowup of the data in (b) between 10 s and 10.5 s. (d) Frequency spectrum of the recording of Channel 1.
with our current fabrication method is the formation of a coplanar fabrication substrate. The key to the integration of a bonding substrate in the isMEA is to create a heterogeneous coplanar structure as the starting fabrication substrate for the via-bonding process [34]. With the FR4-based PCB as the bonding substrate, the protruding metal structures ( high) caused a device yield problem (the current yield is between 50~90%). On the one hand, metal contact pads with such a height constrained the microfabrication process to only form shallow inclined-vias on them; on the other hand, the metal contact pads had rough surfaces and edges which could interfere with the photolithography process to form badly-shaped inclined-vias through the PDMS overlay. These two issues led to a higher failure rate for inclined-via based interconnects patterned on such FR4-based PCBs. To mitigate this problem, we propose to use flexible PCBs with thinner metal traces. In a pilot experiment, we used a patterned copper-on-PET (Polyethylene terephthalate) sheet (18 copper foil laminated on a 25 PET film) as the bonding substrate and the inclined-via
fabrication was encouraging. In contrast, when fabricating the IC-integrated isMEA shown in Fig. 5, we had no such issue on heterogeneous coplane formation. B. Device Stretchability Our previous interconnect stretching study showed that the gold interconnects could withstand longitudinal tensile strains in the range of 1 ~ 2% [33]. Such stretchability is considerably lower compared to that reported of flat traces produced by other groups [43], [45]. This discrepancy may be the result of different micro/nano-scale gold film morphologies deposited with different equipment under different conditions [40]. Another possibility is that our tests did not adequately assess the stretchability of the interconnects. It was possible that the junction between the trace and the contact pad (see [33, Fig. S3c]) was the site where excess strains generated and broke the trace first when the trace was subject to a strain loading. This would consequently obscure the experimental quantification of the real interconnect stretchability. Therefore, our measurement might
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have underestimated the interconnect stretchability. Further investigation is needed to determine the stretchability of our sandwiched interconnects more accurately. Our finite element modeling of the isMEA’s mechanical performance in epimysial interfacing conditions revealed intuitively how the strain generated in the interfaced substrate was transferred into the gold interconnects and provided insights into layout of the interconnects during isMEA designing. However, it is worth noting that there were two major limitations of this model. First, perfect bonding was assumed between the isMEA and the stretching substrate, whereas in an epimysial interfacing condition, a physiological solution layer exists between the isMEA and the muscle surface. This solution layer acts as lubricant to reduce the degree of strain that can be effectively transferred to the isMEA, so that the actual strains induced in the gold interconnects should be smaller than the simulation results. Second, loading conditions below the fracture limit cause micro cracks to develop in the gold film in response to elongation of the elastomeric substrate. These micro cracks release the induced stress and allow the traces to maintain electrical conductivity in the form of a percolating network [46]. This mechanism, however, was not implemented in our finite element model. In actual gold interconnects under strain loading conditions, when excess strains develop at the characteristic sites [see Fig. 8(a) and (b)], micro cracks are probably generated to release the embedded stress in the gold film, and the interconnects would still maintain their electrical conductivity as a whole, albeit that the interconnect resistance will increase. Fortunately, it was reported that a stretchable MEA could maintain relatively stable electrode impedances even with increase of the interconnect resistance under uniaxial strain loadings [47]. Therefore, the actual strain induced on the gold interconnects of our isMEA under epimysial interfacing conditions can be substantially lower than the fracture limit that completely breaks the interconnects, and may have little effect over the electrode impedance and EMG signal fidelity. In addition, the induced strain on the gold interconnects can be minimized by positioning the conducting layer at or as close as possible to the neutral plane of the sandwiched stretchable MEA structure [48], [49]. These analyses thus help to validate the isMEA’s capability as an epimysial interface, and, indeed, the isMEA functioned properly in our epimysial recording experiments. Previous study on stretchable electronics with glued components revealed that the soft–rigid material interface is the most vulnerable location for causing mechanical failure during deformation [45]. In our via-bonding technology, the bonding between the PDMS overlay and the rigid substrate is usually strong, and it can be further improved or strengthened by brief oxygen plasma treatment of the rigid substrate before applying PDMS coating. Therefore, via-bonds on the rigid components are expected to be strong enough to withstand a significantly large amount of stress [34]. Furthermore, in the isMEA design, the rigid component area was reinforced with additional PDMS. Chances for this rigid component area to experience higher strain than the isMEA body are very small, thus this area should not be the concern for causing device failure during deformation.
C. Hermetic Packaging Our isMEA technology is promising for chronic applications. However, it is known that PDMS is not a hermetic material for electronic packaging. For simplicity and based on our current acute application, in the two isMEA designs presented in this paper, we didn’t take efforts to address this hermetic packaging issue. For chronic implantation applications, a feasible solution is to first encapsulate the entire rigid component (e.g., the PCB or the silicon die) with a fine and stable polymer (such as parylene or polyimide) coating, while only exposing the contact pads, and then start the isMEA fabrication [50]. Passivation using parylene or polyimide is simple and easy using standard microfabrication techniques. Therefore, our isMEA technology is applicable to chronic implantation applications. D. Further Characterization In this paper, we primarily characterized the isMEA in the context of acute epimysial interfacing, because of our current research focus on muscle electrophysiology. The biocompatibility testing in the current study was very preliminary. Further experiments are needed to fully evaluate the long-term in vivo performance of this isMEA technology, including biocompatibility, durability and functional stability. E. Potential Applications This PDMS-based isMEA technology may lead to promising neural interfacing applications that require conformable device contact with biological tissue surfaces and minimal tissue damage. We have demonstrated the epimysial recording capability of an isMEA in this paper, and in a previous study [51], we presented some preliminary results on spatiotemporal epimysial stimulation of a muscle via an isMEA. While our current application focus is on epimysial recording and stimulation, this isMEA technology is also being adapted to peripheral nerve interfacing [36]. Other attractive applications comprise high-resolution retinal prostheses [24]–[26], spinal-cord surface stimulation for prosthetics [17], [38], [52], and high-resolution ECoG-based BCIs [21]–[23]. V. CONCLUSIONS In this paper, we described a PDMS-based isMEA technology that addressed the challenge of effectively packaging stretchable MEAs for connection to standard rigid electronics. Combined with a previously developed high-density interconnect technology, this isMEA technology facilitates a high-resolution, high-density integrated system solution for neural and muscular surface interfacing. This isMEA technology is particularly promising for implanted neural surface interfacing applications that demand high-density microelectrodes and integrated electronics, such as retinal prostheses and ECoG-based BCIs. In the future, we will further evaluate and characterize our isMEA technology in chronic implantation studies, and we will also explore other exciting applications for this isMEA technology (e.g., peripheral nerve interfacing and high-resolution ECoG-based BCIs).
GUO et al.: A PDMS-BASED isMEA FOR NEURAL AND MUSCULAR SURFACE INTERFACING
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Gareth S. Guvanasen (S’05) received the B.S. and M.S. degrees in electrical and computer engineering from Duke University, Durham, NC, and the Georgia Institute of Technology, Atlanta, in 2008 and 2010, respectively. He is currently pursuing the Ph.D. degree in electrical and computer engineering at the Georgia Institute of Technology. He has held internships at ENSCO Inc., Springfield, VA, and Duke Robotics, Durham, NC, during which time he helped design autonomous ground and underwater vehicles. His research interests include functional electrical stimulation, neuroprosthetics, and control systems.
Xi Liu received the B.E. and M.E. degrees in aerospace engineering from the Beijing University of Aeronautics and Astronautics, Beijing, China, in 2003 and 2005, respectively, and the M. S. degree in mechanical engineering from the Georgia Institute of Technology, Atlanta, in 2010. He is currently pursuing the Ph.D. degree in mechanical engineering at the Georgia Institute of Technology, where he is a Research Assistant in the Department of Mechanical Engineering. His research interests include reliability analysis of 3D electronic packages integrated with TSVs and microbumps, and micro-strain measurements.
Christopher Tuthill received the B.S. degree in electrical and biomedical engineering from the Illinois Institute of Technology, Chicago, IL, in 2006. He is currently pursuing the Ph.D. degree in biomedical engineering at the Georgia Institute of Technology, Atlanta. His research interests include motor control, neural control of prosthetic devices, and rehabilitation technology.
T. Richard Nichols received the B.S. degree in biology from Brown University, Providence, RI, in 1969 and the Ph.D. degree in physiology from Harvard University, Cambridge, MA, in 1974. He was a member of the Department of Physiology at Emory University, Atlanta, GA, from 1983 until 2007, after which he became Professor and Chair of the School of Applied Physiology at the Georgia Institute of Technology, Atlanta. His research interests include the neural regulation of limb mechanics, proprioception, locomotion, spinal cord, and peripheral nerve injury.
Stephen P. DeWeerth (S’85–M’90–SM’03) received the M.S. degree in computer science and the Ph.D. degree in computation and neural systems from the California Institute of Technology, Pasadena, in 1987 and 1991, respectively. He is currently serving as Professor and Founding Chair of the Department of Biomedical Engineering at Khalifa University, Abu Dhabi, United Arab Emirates. He is also a Professor in the Wallace H. Coulter Department of Biomedical Engineering, Georgia Institute of Technology and the Emory University School of Medicine, Atlanta, GA. His research focuses on the implementation of neuromorphic electronic and robotic systems, the development of neural interfacing technologies, and the study of the biological control of movement.
