Integration of High-Charge-Injection-Capacity Electrodes onto

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Integration of high charge injection capacity electrodes onto polymer softening neural interfaces David Eduardo Arreaga-Salas, Adrian E. Avendano-Bolivar, Dustin Simon, Radu Reit, Aldo Garcia-Sandoval, Robert Rennaker, and Walter E Voit ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.5b08139 • Publication Date (Web): 17 Nov 2015 Downloaded from http://pubs.acs.org on November 24, 2015

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ACS Applied Materials & Interfaces

Integration of high charge injection capacity electrodes onto polymer softening neural interfaces David E. Arreaga-Salas ¥, Adrian Avendaño-Bolívar, Dustin Simon, Radu Reit, Aldo GarciaSandoval, Robert L. Rennaker, Walter Voit*. Department of Materials Science and Engineering ¥, Department of Bioengineering ϯ and Department of Mechanical Engineering‡. The University of Texas at Dallas, 800 W. Campbell Rd. Richardson, Texas 75080-3021 KEYWORDS: Softening neural interfaces, microfabrication, electrical aging, iridium electrodes, cyclic voltammetry, electrochemical impedance spectroscopy

ABSTRACT

Softening neural interfaces are implanted stiff to enable precise insertion and soften in physiological conditions to minimize modulus mismatch with tissue. In this work, a high chargeinjection capacity iridium electrode fabrication process is detailed. For the first time, this process enables integration of iridium electrodes onto softening substrates using photolithography to define all features in the device. Importantly, no electroplated layers are utilized leading to a highly scalable method for consistent device fabrication. The iridium electrode is metallically bonded to the gold conductor layer, which is covalently bonded to the softening substrate via sulfur-based

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click chemistry. The resulting shape memory polymer neural interfaces can deliver more than 2 billion symmetric biphasic pulses (100µs/phase), with a charge of 200 µC/cm2 and geometric surface area (GSA) of 300 µm2. A transfer-by-polymerization method is used in combination with standard semiconductor processing techniques to fabricate functional neural probes onto a thiolene based, thin film substrate. Electrical stability is tested under simulated physiological conditions in an accelerated electrical aging paradigm with periodic measurement of electrochemical impedance spectra (EIS) and charge storage capacity (CSC) at various intervals. Electrochemical characterization, and both optical and scanning electron microscopy suggest significant breakdown of the 600 nm thick parylene-C insulation, although no delamination of the conductors or of the final electrode interface was observed.

Minor cracking at the edges of the thin film

iridium electrodes was occasionally observed. The resulting devices will provide electrical recording and stimulation of the nervous system to better understand neural wiring and timing, target treatments for debilitating diseases and give neuroscientists spatially selective and specific tools to interact with the body. This approach has uses for cochlear implants, nerve cuff electrodes, penetrating cortical probes, spinal stimulators, blanket electrodes for the gut, stomach and visceral organs and a host of other custom nerve-interfacing devices.

INTRODUCTION The neuroscience community faces a fundamental challenge limiting the widespread use of neural interfaces for diagnostics; research or treatment: the lack of a chronically stable connection between synthetic devices and the nervous system. Over the past several decades, the community has made remarkable advances from initial studies that inserted conductive wires into frogs more than 100 years ago 1 to recent work that enables quadriplegic patients to volitionally control robotic arms for short periods of time.2 Despite the remarkable developments


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in the design and fabrication of reliable neural interfaces, the hard-to-avoid immunological reaction from neurological tissue to foreign devices has limited chronic functionality. 3-4 Many different approaches have been proposed to improve the biotic-abiotic interaction and the longevity of the implant. Miniaturization of probes has been studied by several groups with the goal of making the foreign device less invasive to the nervous system. In 2012 Kipke et al. introduced an electrode 10-times smaller than conventional devices. These probes, with a carbon fiber core and a poly(3,4-ethylenedioxythiophene) (PEDOT-PSS) electrode interface, were capable of recording single-unit activity acutely in rats. The so-called ultra-small devices elicited a reduced immune response compared to that of silicon probes. 5 More recently Anikeeva and colleagues developed a bifunctional polymer-fiber probe fabricated through a thermal drawing process. 6 This single-channel device was capable of recording neural signals, using doped carbon black coatings at the tip and stimulating optically by guiding light throughout a polycarbonate core. This mechanically-compliant device was designed to withstand the movements of the limb and other high-strain applications. Conventional devices have been coated with different materials to improve their biocompatibility. Tresco et al. incorporated permeation layers onto conventional silicon-based devices to create preferential diffusion of proinflammatory molecules into the substrate to minimize tissue response. It was demonstrated that over a 16-week span, the foreign-body response to coated with a permeable layer was less compared to uncoated probes. 7 Under a different approach, more compliant materials have been used to fabricate flexible neural interfaces to reduce mechanical mismatch between neural tissue and the foreign implanted device.9-11 Kim et al. built a three-dimensional parylene-C probe with a geometric design that matched the anatomy of the barrel cortex. These devices were loaded


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with different anti-inflammatory drugs to reduce the immune response. However given the mechanical compliance of these devices, external insertion tools are required. 12 The hypothesis that softer, modulus-matched devices may fare better in nerve tissue is not new. In 2005, Kipke et al. used finite element modeling to assess the potential benefits of a “hypothetical soft probe” of near 6 MPa to reduce the tangential stress between the device and neural tissue. 13 The first attempt to demonstrate this hypothesis was performed using bioinspired composites able to alter their modulus gradually upon exposure to physiological conditions. 14 Capadona et al. implanted intracortical passive probes made out of the moduluschanging composite to analyze the temporal tissue-response at different periods of time and compared immunological reactions versus rigid probes. Neuron population and leakage of the brain blood barriers (BBB) did not show a statistically significant difference between the two types of probes after 8 weeks. However after 16 weeks a larger neuron density and improved BBB stability was observed around the mechanically adaptive probe. 15 16 Following that pioneering experimental work, shape memory polymers (SMP) were explored as a different approach for mechanically-adaptive probes. 17 In 2012, acrylate systems were chemically tuned to trigger a decrease in modulus from 700 MPa down to 300 kPa when exposed to physiological conditions. However, this SMP system has limited compatibility with standard microfabrication techniques. Later a thiol-ene/acrylate system with improved thermomechanical stability was introduced. Microelectrode arrays with small geometric surface area (GSA) electrodes were fabricated using standard photolithographic techniques. Dimensional stability above the glass transition temperature allowed the use of standard semiconductor fabrication tools in large area samples without alignment limitations; a key factor for scalability. 18 That work demonstrated for the first time, a polymer system that could function as a substrate for precision photolithographic


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processing and thin film fabrication at temperatures above the glass transition temperature of the polymer. This milestone represented a paradigm shift not only for bioelectronic applications but also for the flexible electronics industry in general, according to issues laid out succinctly by Wong et al.19 All the previously mentioned softening devices have lacked an electrode-interface material that provides an electrochemically reliable interface to neural tissue. When microfabrication processing compatibility is jeopardize by the substrate, electroplated materials appear as viable option to improve the electrochemical interaction properties of implantable electrodes. PEDOTPSS, 20-25 electroplated iridium oxide (IrOx) 26-27 and carbon nanomaterials 28-33 have been used as methods to reduce interface impedance and to increase CSC. Carbon nanomaterials achieve this mainly by increasing the electrochemical surface area (ESA), while PEDOT-PSS and IrOx introduces a dominant faradaic charge transfer mechanism. However, premature delamination and cracking during prolonged studies as well as large variability between samples has been reported repeatedly in literature. 21, 34 A widely accepted electrode material for chronic recording and stimulation applications in implantable devices is iridium oxide films: either sputtered iridium oxide films (SIROF) or activated iridium oxide films (AIROF). 35-40 Cogan et al. have studied the properties of AIROF and SIROF in detail by electrochemical impedance spectroscopy (EIS) and cyclic voltammetry (CV). AIROF electrodes are fabricated on iridium metal by cyclic oxidation and reduction of the metal surface after the electrode array has been fabricated. This activation process can be done in saline solution or in vivo. 41 SIROF is sputtered from a pure iridium target in an oxygenated atmosphere. The goal of this work is to develop a fabrication process to incorporate reliable high performance electrodes onto thiol-ene based softening substrates using standard semiconductor


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fabrication techniques. This is a critical step in the evaluation of the potential benefits of using mechanically adaptive neural interfaces in chronic studies and eventually for clinical use. Materials selection is of high importance for the integration of electronic elements onto a flexible substrate. In the specific case of softening neural interfaces, thin film stress, interlayer adhesion, etching selectivity and strain tolerance are critical factors in the selection of materials and processing techniques. The coefficient of thermal expansion (CTE) mismatch between iridium and the underlying materials requires the employment of different strategies to confine and mitigate the thin film stresses. 42 Interlayer adhesion, a key factor in the robustness of devices, is higly dependant on the deposition and polymerization conditions. 18, 43 The difficulty in etching iridium metal constrains the ability to pattern it, leading to the use of lift-off techniques. 24, 44 For the interaction of devices with the peripheral nervous system, strain tolerance is also a consideration. While iridium excels as an electrochemical interface and as a conductor, it does not possess the ductility required to serve as a conductor for strain demanding applications. 45-47 Therefore, processing methods such as the ones presented in this work are important to enable multilayer devices with high channel density, strain tolerant conductors, and encapsulation layers independently patterned onto compliant substrates. For this to be possible, the underlying substrate has to be able to withstand the thermal cycles that photolithography implies, without shrinking or expanding permanently. 42 In this work a comprehensive fabrication protocol for iridium-functionalized softening neural interfaces is developed. Thin film stress confinement and interfacial stress passivation techniques are utilized to enable devices with appropriate mechanical robustness. Electrical chronic functionality of the devices is evaluated through electrical aging in saline solution to understand


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possible failure mechanisms and set the basis for future generations of functional mechanically adaptive neural interfaces.

RESULTS A novel process to fabricate robust iridium electrodes onto physiologically responsive substrates.is presented in this work This is accomplished through the use of a transfer-bypolymerization technique 17, 48 with a modified device stack to enable a highly scalable, robust process to fabricate flexible electronics and specifically neural bioelectronics. The biological response of softening shape memory polymer based cortical probes was compared to high modulus parylene-C based probes. The electrochemical properties of the iridium electrodes were assessed via microscopy, thermomechanical analysis and electrochemical analysis during a 169 hour test at various points during which more than 2.2 billion stimulation pulses were passed through the iridium electrodes. Specifically, scanning electron microscopy (SEM), optical microscopy, dynamic mechanical analysis (DMA), cyclic voltammetry (CV) and impedance spectroscopy (IS) are utilized to measure the response of devices to simulated physiological conditions. Figure 1(a) places the described devices in context to other probes, devices and cortical tissue according to the Young’s modulus of the material. Immunohistochemistry comparison of smart softening and parylene-c based probes. Figure 1(b) presents an immunohistochemical comparison between a parylene-C probe, with a modulus of 4.5 GPa at 38 ⁰C, and a physiologically responsive softening device, that lowers its modulus from 700 MPa to nearly 40 MPa after exposure to physiological conditions.48-49 Blood brain barrier (BBB) disruption is quantified by staining 30 µm thick microtomed sections of adult male Sprague Dawley rat cortical tissue one month after the surgical procedure to implant


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the polymeric devices made from parylene-C and softening shape memory polymer. The immunoglobulin G (IgG) stain helps identify the presence of IgG in the cortical tissue which is not normal in unperturbed conditions, and typically correlated to the leakage of blood-derived cells and serum proteins. 50-51 Figure 1(b) is a representative micrograph of a tissue slide stained for IgG from one of the five animals that were dually implanted with parylene-C and thiol-ene based intracortical probes. IgG intensity is proportional to brain-blood barrier rupture after four weeks of implantation. In general a larger IgG intensity is observed around the parylene-C intracortical implant near the voids left after probe explantation. Figure 1(c) is the quantified IgG response across all five animals versus the distance form implant. At distances of up to 80 µm from the implant, significant differences in IgG response are observed between the paryleneC and softening thiol-ene based substrates. Fabrication of functionalized softening micro-electrodes The key steps in the fabrication of functionalized softening microelectrodes are presented in Figure 2 (detailed steps can be found in the experimental section). Iridium electrodes of 300 µm2 GSA were patterned on a sacrificial glass substrate. In order to have a high yield during the transfer-by-polymerization process the polymer substrate thickness was kept below 50 nm. Due to the strong chemical interaction between thiol-ene based polymers and gold, a 300 nm gold layer was used as an adhesion interface layer between the iridium electrode and the responsive polymeric substrate. The gold layer was also used as the conductor material to electrically connect the final electrode interface and the recording/stimulation system. Parylene-C was used as electrical insulation layer. Parylene-C is widely used in flexible electronics because of its room-temperature deposition capability, and its relatively low moisture absorption. All devices


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in the study were fabricated with 4 shanks, and 4 channels on each shank, as presented in the optical micrograph shown in Figure 2. Electrical aging and failure mechanism Iridium microelectrodes transferred onto the softening thiol-ene/acrylate substrates were electrically aged using a 200 µC/cm2 pulse at a rate of 3600 Hz. Figure 3 is a schematic that represents the experimental setup. Full details of the electrochemical characterization can be found in the experimental section. Notable for understanding and interpreting data in the remainder of this section are the precision with which the reference (Ag|AgCl) electrode, the counter (platinum wire) and the working electrodes (the device) are placed within the phosphate buffered saline (PBS) bath at 37 °C relative to one another to achieve consistent results. The optical micrograph in Figure 3 pictures the full connected device from which the shanks in Figure 2 were observed. Iridium electrodes studied in this work were not deliberately activated nor was further understanding of the film’s stoichiometry performed. Figure 4 describes the results of high duty cycle electrical aging of devices, encompassing approximately 2.2 billion biphasic pulses of stimulation over 169 hours. Figure 4 (a) shows the measured transient voltage curves (without iR compensation) of a representative channel captured at different time points during the length of the experiment with initial results in cyan and results after 169 hours in red. The intermediate transient voltage curves were measured at time points after the first 2, 8, and 16 hours and then approximately every 12 hours after that. Light gray dotted lines and the initial and final lines represent averages of five channels across the 16 electrodes. Four channels were not cyclically stimulated to be used as control reference channels and seven channels were stimulated but not measured at each time point as a second kind of control.


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The bottom section of Figure 4 (a) represents the geometry of the charge-balanced biphasic input current pulse used during the experiment. The positive and negative regions of the driven symmetric square wave current each lasted 100 µs and were separated by a 25 µs interface time (after the negatively driven current) and a 75 µs idle time between neighboring biphasic pulse trains. Gradually, throughout the experiment the potential required to drive a fixed charge density and maintain the same biphasic square wave current pulse trains decreased. The maximum negative potential excursion (Emc)52 decreased nearly 0.1 V while the access potential (Va) remained roughly the same (0.12 V). Figure 4 (b) presents the average EIS of 5 different active electrodes at different time points. In the high frequency range (near 10 kHz) the electrochemical impedance gradually and consistently increases over time from approximately 10 kΩ up to 25 kΩ (consistent with the transient voltage in Figure 4(a)). The impedance value does not present a significant change near the physiologically relevant frequencies (1 kHz – 4 kHz). At lower frequencies—100 Hz and below—the impedance decreases compared to the values recorded before electrical aging. The phase angle shifts at 100 kHz from -45 ° (initial conditions) up to -70 ° (after 2.2 billion stimulation pulses during the 169 hour test). Also there is a peak initially located at 10 kHz that gradually shifted towards the 100 kHz region. At the lower frequencies (1 kHz), the phase angle shifted from -75 ° to -55 °. When approaching direct current the phase angle moves from -75 ° to -45 °. Representative curves of cyclic voltammetry from one representative electrode channel are shown in Figure 5 (a), three curves are highlighted. (Initial, after 49 and after 92 hours). In Figure 5 (b) the Young’s modulus (measured in compression via a rheometer) of the smart polymer substrates are plotted (left axis) versus swelling time. The average Young’s modulus is


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presented at each time point: Dry, 1, 4, 14, 24, 48 and 96 hours of swelling. After 48 hours the modulus reaches a plateau at 36 MPa. In the same figure but on the right axis, average cathodic charge storage capacity (CSCc) is presented versus time (under electrical stress). No significant change is observed before 27 hours; however after 27 hours, CSCc begins to increase from below 10 mC/cm2 to more than 70 mC/cm2. The passive electrode exhibits an insignificant lower average CSCc at longer time points. Optical and electronic microscopy of post aging experiments are presented in Figure 6. Residues of the phosphate buffer solution (PBS) can be observed on top of the encapsulation and the electrode pads in both micrographs. Cracks can be appreciated on the parylene-C encapsulation at different places, but no delamination of the iridium layer was observed in or on the electrode pads. DISCUSSION Not only have we shown a novel process to fabricate the first thin film iridium electrodes made on physiologically responsive polymers, we have demonstrated survivability in high duty cycle electrical aging conditions on softening substrates that have been shown to reduce the chronic physiological response in one month rat studies. This manuscript lays the groundwork for a class of high charge-injection capacity neural microstimulators with the potential to help understand disease biology, map neural circuitry and modulate the body’s nervous system. Notably, after more than 2 billion stimulation pulses, the iridium and shape memory polymer portions of the electrode remained intact while the thin parylene-C coating exhibits cracking leading to shunting, and an increase in the effective surface area of the electrode. The rationale for the length and conditions of the aging experiment are based on of the most widely used neural interfaces by the medical community, cochlear implants (CI). CI typically


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last between 5 to 10 years before a re-implantation is required. 53-54 Implants stimulate for on average 14 hours per day at a mean charge density of 25 µC/cm2 based on the characteristics of the environment in which patients live, 55 at a typical frequency of 100 Hz. 56 This is roughly 16 billion pulses over the pass of 5 years. 2 billion stimulation pulses at 8× the charge density (200 µC/cm2) was chosen as an approximate accelerated electrical aging to the functional lifetime of a CI. While many other variables are at work and simulated environments do not fully represent long term chronic implants, the early indications for these devices, and for next generation devices with improved encapsulation are quite promising. The initial rationale of this work lies in the need of functional electrode interfaces to be incorporated robustly onto stimuli-responsive devices. This will allow neuroscientists and someday clinicians to take full advantage of the improved biological response vis-à-vis reduction of mechanical mismatch with tissue of softening devices as suggested by Kipke et al. 13 The brief immunohistochemistry analysis presented in Figure 1(b) demonstrates the potential of using thiol-ene based softening devices to reduce the disruption of the BBB over extended periods of time. These results are consistent with the results presented previously for other softening systems. 16-17, 57 In Figure 1(c), quantified IgG intensity across five implanted animals indicates that the reduction in mechanical modulus brings a significant IgG reduction near 80 µm from the implant. This trend becomes more obvious approaching 50 and 25 µm from the penetration site. To fully enable this technology in chronic studies and subsequent clinical use, electrical reliability is of major relevance. Throughout this work, further evidence of the potential impact of using softening neural interfaces is collected. A comparison was made of devices made with identical design and dimensions and fabricated on flexible substrates, both non-softening and softening. Because inflammatory events, resulting from neurological disease or device


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implantation, can cause the BBB to remain open, 50 it is hypothesized that the reduced modulus of the softening substrates chronically allows a better resealing of the BBB and reduces the number of inflammatory events. This hypothesis that softening substrates reduce the inflammatory response during chronic penetration through the BBB has not been conclusively proven and remains an open problem in the literature and is ripe for future studies, although our results give indication that this may be the case. The chosen device stack is important. To achieve a strong adhesion between the final electrode interface and the conductive layer, direct deposition of iridium onto gold layers has been shown to be effective. 43, 58 Nevertheless the fabrication of softening-iridium devices using a bottom-up fabrication approach is limited by the poor etching selectivity of iridium and the high thin film stress of iridium when deposited directly on the softening substrate. Therefore the development of a modular, full photolithography fabrication method to integrate iridium electrodes in softening devices becomes of high importance. The fabrication process described in Figure 2, allows the integration of partially confined iridium features in an island structure onto smart softening substrates. Transferring small isolated features from a sacrificial glass carrier, helps reduce the interfacial stresses that could be generated when deposited directly onto the flexible substrate. However due to the nature of the transfer-by-polymerization method, adhesion forces define the effectiveness of the process. In the case of the stack glass/iridium/gold, only thin (> 50 nm thick) iridium features can be transferred reliably with a high yield. The optical micrograph shown in Figure 2 presents a fully processed device, which after several thermal cycles—typical during photolithography—remains structurally intact. This indicates that the island-like structure reduces the stresses across the stack of the device. In failed experiments, thicker iridium samples and much larger unpatterned


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iridium surface cracked due to the buildup of stresses on the thin film interface during the transfer process. This sets some boundary conditions on the broad applicability of our approach, but could be overcome with more complex multi-layer device stacks to distribute stresses across several layers. It is necessary to understand the mechanical and electrical stability of the electrode materials at the device level under simulated physiological conditions and depict possible failure mechanisms. The length of the experiment and the charge utilized were selected based on the 5 year life-time projection of one of the most demanding, clinically used neural interfaces, a cochlear implant. . In order to isolate the effects of the electrical aging from the mechanical changes in the softening substrate, 4 electrodes were only exposed to the temperature and humidity conditions. With the purpose of maintaining the sample unperturbed during the length of the experiment, the electrical setup was configured to switch between the current train generator (Plexon Inc. Stimulator 2.0) and multiplexed electrochemical characterization station (CH Instruments Inc. CHI660D/CHI684). Oxygen was removed from PBS solution by gentle argon gas flow. pH was monitored during the length of the experiment with an external bench to pH meter (Mettler Toledo Inc.). A detailed understanding of the electrochemical setup is critical for proper interpretation of data. In the voltage excursions presented in Figure 4 (a), there is a gradual reduction of the maximum voltage (Emc); however the access voltage (Va) remains practically unchanged at -0.12 V. Va is associated with the ohmic resistance of the electrolyte and concentration polarization, from the maximum negative voltage in the transient. Emc meanwhile is a measure of the polarization across the electrode-electrolyte interface. 41 Therefore the drop of 0.1 V in Emc represents an increase in charge injection capacity that can be attributed to the initial


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electrochemical activation of the iridium. EIS in Figure 4 (b) allows observation of an interesting behavior in the evolution of impedance spectra and the phase angle: around 1 kHz the impedance of the electrode after 169 hours is relatively lower and the phase angle indicates a dominant faradaic reaction—typical of iridium electrodes.52, 59-60 When exploring low frequencies, the impedance is nearly a full order of magnitude lower. Although a lower impedance is expected due to the activation of iridium, such trend suggests that the electrode surface is not the only changing element in the device. Cyclic voltammetry was also performed during the length of the experiment at different time points. In Figure 5 (a), three representative curves are shown: initial electrical observations, observations after 49 and observations after 92 hours. Between the initial voltammogram and the one collected after 49 hours there is a slight increase in CSC; this increase is further accentuated after 92 hours and remains unchanged after that. The increase in CSC could be related to the activation of the iridium electrodes; however, comparison of the CSCc of active and passive electrodes suggests an alternative explanation for the increase in CSC. Figure 5 (b) presents in the left axis the Young’s modulus progression of the softening substrate over soaking time. It is important to mention that the largest drop in modulus occurs during the first 48 hours of plasticization, after which a plateau of roughly 40 MPa is reached. Interestingly, on the right axis the CSCc of the passive electrodes also raises near the 40 hours range, without an electrical stimulation that could have activated the electrodes. This change is an indication of a physical change, and more specifically a change in the electrical insulation. In Figure 6, post-aging optical and electron microscopy reveal multiple cracks in the parylene-C encapsulation, something previously seen in literature.11, 61-63 The newly available electrochemical surface area could interact with the electrolyte and increase the CSCc in both passive and active devices. These


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changes are observable in the CVs due to the slow scan rate at which they were performed. At 50 mV/s, ions are able to move through the cracks in the encapsulation and interact with gold traces; this is not possible at high frequencies (> 1 kHz). Iridium electrode pads remain physically intact, with no observable changes that would indicate delamination of the island-structured feature. It is hypothesized that the change in modulus of the softening substrate introduces larger interfacial stresses between parylene-C and the softening substrate. These interfacial stresses, in addition to weak chemical adhesion, leads to the failure of the encapsulation and subsequently to electrical instability at low frequencies. Experimenting with different levels of softening and softening encapsulation layers are active areas of research that could overcome these challenges. CONCLUSION A novel fabrication method to integrate iridium metal electrodes onto smart softening flexible substrates has been demonstrated. Its relevance lies in the exclusive use of standard microfabrication techniques that can enable high yield, large scale and repeatable fabrication processes. Iridium softening-devices demonstrated long-term functionality in saline solutionand withstood over two billion stimulation cycles. Transient voltage degradation was not observed during the length of the experiment. Monitored impedance values at low frequencies and CSC calculated form CV at slow scan rates, suggest the emergence of newly available electrochemical surface area. Imaging confirms that the increase in electrochemical surface area, is due to encapsulant (parylene-C) cracking. However the underlying stack did not suffer of delamination or cracking. These are very encouraging results that motivate the research of more compliant encapsulation layer with improved barrier properties that enable the advancement of mechanically adaptive neural interfaces.


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EXPERIMENTAL SECTION Immunohistochemistry: All animal testing was conducted in accordance with the University of Texas at Dallas Institutional Animal Care and Use Committee (IACUC) regulations. Surgical procedure, slide preparation methods and optical micrographs of the devices implanted are included within the supporting information document. The devices utilized consisted of four 1.5 mm long shanks that were 150 μm wide and 35 μm thick. Five animals were implanted at a depth of approximately 500 μm. A profiling method was developed wherein the user selects the two end points of the track along the center of its longest axis, and the program measures 50 evenly spaced intensity profiles perpendicular to this centerline in each direction. Images are converted to greyscale, then threshold-based analysis is used to select the dark area of the track void, which is excluded from the intensity profile analysis in order to control for variations in void size and shape, as well as better characterizing the device-tissue interface. The selected track void area is overlaid with a black mask, and the track centerline is then selected manually. Each profile line consists of the first 200µm of pixels with a value >0 (e.g. beyond the void area mask). The resulting 100 profiles are averaged together to yield a mean profile for each track, which is then separated into 25µm bins for analysis. Intensity values shown represent mean pixel values for each 25µm bin. To control for animal-to-animal variations in native expression levels of the markers of interest, intensity measurements were taken from the non-implanted contralateral hemisphere and the mean values for each rat were subtracted from the corresponding mean implant profiles. Background intensity profiles were taken from the same locations in the image frame as the experimental tracks in order to account for spatial non-uniformities in microscope illumination.64


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Device Fabrication The fabrication process steps are here presented in the order they were performed and referred to in Figure 2. Sacrificial carrier preparation: Glass microscope slides of 75×50 mm were thoroughly cleaned. All optically visible undesirable material was removed by scrubbing in an Alconox solution, sonication in acetone, sonication in isopropanol and repeated as necessary. Iridium electrodes deposition patterning: To outline the iridium electrodes a 1.3 µm layer of photoresist (S-1813 from Microposit Inc.) was spin-coated on a sacrificial substrate. Using a Karl Suss contact printer (MAB6) and manufacturer recommended parameters the photoresist was UV exposed and chemically developed. Iridium pellets (99.99%) were purchased from Kurt J. Lesker Company. Iridium metal was deposited on top of the patterned sacrificial lift-off layer by electron beam evaporation at a 0.2 Å/sec rate. Unwanted iridium was removed by sonication in acetone for 10 min. Samples were subsequently rinsed in isopropanol to remove acetone residues. Metal conductor deposition: A 300 nm of gold blanket was deposited on a iridiumpatterned/sacrificial carrier at 2 Å/sec, using electron-beam evaporation. Monomer solution preparation: Tricyclodecane dimethanol diacrylate (TCMDA), 1,3,5-triallyl1,3,5-triazine-2,4,6(1H,3H,5H)-trione (TATATO) and 2,2-dimethoxy-2-phenylacetophenone (DMPA) were purchased from Sigma Aldrich and used as received. Tris[2-(3mercaptopropionyloxy)ethyl] isocyanurate (TMICN) was purchased from Wako Chemicals. The monomer solution was prepared using 31 mol % of TCMDA, 34.5 mol % of TATATO, 34.5 mol % of TMICN and 0.1 wt% of DMPA.


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Transfer-by-Polymerization: Monomer solution was cast on gold/iridium/sacrificial-carrier stack. Paper spacers were used to set the thickness at 35µm and a clean glass slide was put on top to serve as carrier for the rest of the process. The mold with the monomer solution was put in a crosslinking chamber with five overhead 365 nm UV bulbs (UVP via Cole-Parmer) for 45 minutes. After polymerization the sacrificial carrier slide was removed leaving a stack of iridium/gold/polymer on a processing carrier. More details of the transferred by polymerization method have been published elsewhere.48 Gold conductor patterning: Transferred gold layer was patterned as the conductor material by standard photolithography and etched using gold etchant (AU-5 Cyantek Corporation). Parylene-C deposition and patterning: As measured by optical ellipsometry on a witness sample, 600 nm of parylene-C were deposited with a SCS Labcoter 2 (SCS Systems).Using oxygen plasma in a reactive ion etching system, the parylene-C layer was patterned to function as electrical insulator. Patterning of softening substrate: Both active and passive (no electronics built-in) devices were patterned to the device design using a high power class 4 laser with a spot size of 10 µm in diameter. Electrical Aging and Characterization Electrochemical setup: 1× phosphate buffer solution with a pH of 7.4 was kept at 37 °C. The solution was deoxygenated with argon gas flow during the entire length of the experiment. Smart softening microelectrodes were electrically connected through a 16-channels zero insertion force connector, soldered to a PCB board from one side and to an edge-card male to the other end. A DB25 connector was adapted to the potentiometer-multiplexer (described below) and connected to a female edge card connector. A Stimulator 2.0 from Plexon Inc. (Dallas TX) was used to


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generate a current-controlled, charge balance, cathodic-first pulse train of 6 µA of 100 µs with interphase time of 25 µs and idling period of 75 µs. The electrical stress was only stopped every 12 to 24 hours to record EIS and CV as described below. Four of the 16 electrodes, one from each shank, were not electrically stressed during the experiment, resulting two sets of electrodes; active and passive. Cyclic voltammetry (CV) and electrochemical impedance spectroscopy (EIS): Devices were electrically characterized using a CHI 660D potentiostat connected to a 32 channel CHI 684 multiplexer from CH instruments in Austin TX. All measurements were performed using a three electrode setup: platinum wire as counter electrode, Ag|AgCl as a reference electrode and the device electrode as the experimental electrode. Potential was cyclically varied from -0.6 V to 0.8 V, at a 50 mV/s scan rate. Cathodic charge storage capacity was calculated from the time integral of the cathodic current during one complete potential cycle during cyclic voltammetry. EIS measurements were carried using no initial potential and a 5mV sinusoidal wave, varying frequency from 100 KHz down to 1 Hz. Initial CV and EIS were recorded prior to any electrical stimulation, immediately after immersion in PBS. No prior activation was performed on the electrodes utilized in these tests. Compression modulus measurements: Four disks measuring 3mm in diameter and 1.2mm in thickness were prepared for each measured time point. After soaking in PBS for the allotted amount of time, samples were then transferred to a TA Instruments Rheometer for compressive testing. Samples were placed on a heated stage (37C) and preloaded to 0.02MPa. Finally, PBS was introduced to the air gap and the temperature was allowed to settle for 10 minutes before measurement began.


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Scanning electron microscopy: Samples were left in a vacuum desiccator for 24 hours to remove water. A thin layer of Au/Pd alloy was sputtered onto the devices to avoid accumulation of a negative charge on the surface of the device. Using a Zeiss supra-40, the electron gun was set to 5 KeV. The stage was tilted 45 degrees and adjusted to a working distance of 4 mm.


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FIGURES Figure 1. (a) A schematic representation of the Young’s modulus change of a softening probe when exposed to physiological conditions. (b) A 4-week immunohistochemistry experiment compares the IgG reaction of parylene-C and a softening probes of identical length, width and thickness (see supporting information). The immunoreaction is quantified and presented in (c).


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Figure 2. The schematic represents the fabrication process for the photolithographic integration of iridium microelectrodes (left); detailed steps are described in the experimental section. An optical microscope image is shown of an as-fabricated, softening polymer device containing gold traces and iridium microelectrodes (right).


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Figure 3. The diagram details the electrochemical setup for characterization and aging of smart softening devices with iridium microelectrodes.


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Figure 4. (a) Voltage transient response to the current pulse represented in the bottom of the panel is captured at different time points. Initial and final incursions are highlighted in green and red respectively. (b) Progression of the electrochemical impedance spectrum and phase angle evolution are recorded during the electrical aging process.


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Figure 5. (a) Cyclic voltammetry curves of selected time points are recorded from active electrodes. (b) On the left axis, Young’s modulus (compression) is plotted against swelling (aging) time. On the right axis, cathodal charge storage capacity of both, active and passive electrodes is plotted against time under electrical stress.


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Figure 6. Optical microscopy (top) and electron scanning microscopy (bottom) picture a softening device after 169 hours of electrical aging, which is more than 2 billion biphasic chargebalanced stimulation pulses of ±6 µA at a high duty cycle.


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ASSOCIATED CONTENT Supporting information. Immunohistochemistry methods and protocols used during this study. This material is available free of charge via the Internet at http://pubs.acs.org. AUTHOR INFORMATION Corresponding Author E-mail: [email protected], Phone: 972-883-5788 ACKNOWLEDGEMENTS Special thanks to Dr. Stuart Cogan for his valuable discussions on the electrochemical phenomena. We appreciate the long discussions and support given by members of the Advanced Polymer Research Laboratory who aided with various experimental tasks and understanding. D.A.S, A.A.B and A.G.S are generously supported by the CONACYT program and the Mexican government. R.R. is generously supported by the NSF GRFP and the Founder’s Fellows Program at UT Dallas. W.V. acknowledges generous support from the DARPA Young Faculty Award (D13AP00049), the Center for Engineering Innovation at UT Dallas and the Texas Medical Devices Center at UT Dallas, funding from the GlaxoSmithKline bioelectronic medicines consortium and gifts from Texas Instruments which all supported these efforts. REFERENCES 1. Wells, D. A., The Science of Common Things: A Familiar Explanation of the First Principles of Physical Science. For Schools, Families, and Young Students. Publisher Ivison, Phinney, Blakeman, 1859, 323 pages page 290. 2. Hochberg, L. R.; Serruya, M. D.; Friehs, G. M.; Mukand, J. A.; Saleh, M.; Caplan, A. H.; Branner, A.; Chen, D.; Penn, R. D.; Donoghue, J. P., Neuronal Ensemble Control of Prosthetic Devices by a Human with Tetraplegia. Nature 2006, 442 (7099), 164-71. 3. Christensen, M. B.; Pearce, S. M.; Ledbetter, N. M.; Warren, D. J.; Clark, G. A.; Tresco, P. A., The Foreign Body Response to the Utah Slant Electrode Array in the Cat Sciatic Nerve. Acta Biomater 2014, 10 (11), 4650-60.


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64. Biran, R.; Martin, D. C.; Tresco, P. A., Neuronal Cell Loss Accompanies the Brain Tissue Response to Chronically Implanted Silicon Microelectrode Arrays. Exp Neurol 2005, 195 (1), 115-26.


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Figure 1. (a) A schematic representation of the Young’s modulus change of a softening probe when exposed to physiological conditions. (b) A 4-week immunohistochemistry experiment compares the IgG reaction of parylene-C and a softening probes of identical length, width and thickness (see supporting information). The immunoreaction is quantified and presented in (c). 213x114mm (150 x 150 DPI)



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Figure 2. The schematic represents the fabrication process for the photolithographic integration of iridium microelectrodes (left); detailed steps are described in the experimental section. An optical microscope image is shown of an as-fabricated, softening polymer device containing gold traces and iridium microelectrodes (right). 129x123mm (150 x 150 DPI)


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Figure 3. The diagram details the electrochemical setup for characterization and aging of smart softening devices with iridium microelectrodes. 135x177mm (150 x 150 DPI)



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Figure 4. (a) Voltage transient response to the current pulse represented in the bottom of the panel is captured at different time points. Initial and final incursions are highlighted in green and red respectively. (b) Progression of the electrochemical impedance spectrum and phase angle evolution are recorded during the electrical aging process. 215x111mm (150 x 150 DPI)


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102x134mm (150 x 150 DPI)



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72x105mm (150 x 150 DPI)


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86x34mm (150 x 150 DPI)


Integration of High-Charge-Injection-Capacity Electrodes onto

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