Fluoropolymer-Based Flexible Neural Prosthetic Electrodes for

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Fluoropolymer-based Flexible Neural Prosthetic Electrodes for Reliable Neural Interfacing Yong Hee Kim, Jongkil Park, Ho Koo, Min Sun Kim, and Sang-Don Jung ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b12364 • Publication Date (Web): 29 Nov 2017 Downloaded from http://pubs.acs.org on December 3, 2017

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ACS Applied Materials & Interfaces

Fluoropolymer-based Flexible Neural Prosthetic Electrodes for Reliable Neural Interfacing Yong Hee Kim,a Jongkil Park, a Ho Koo,b Min Sun Kim,b Sang-Don Jung a* a

Synaptic Devices Research Section, Electronics and Telecommunications Research Institute,

218 Gajeong-ro, Yuseng-gu, Daejeon 34129, Republic of Korea. b

Department of Physiology, Wonkwang University School of Medicine, 895 Munwang-ro, Iksan,

Jeollabuk-do 570-711, Republic of Korea. *

Corresponding Author Email: [email protected]

KEYWORDS: fluoropolymers, RF plasma treatment, co-adhesion, epidural electrode, reliability, neural interface

ABSTRACT: We covalently bound fluoropolymer (FP) films by plasma treatment followed by thermal pressing at temperatures below their melting point, and fabricated an adhesion-metalfree flexible gold electrode array entirely encapsulated by the FP film, excepting the active electrode openings. The fabricated device was chemically resistant and was modified to have a lower impedance and efficient charge injection capability. The fabricated device was evaluated in vivo in rats and was confirmed to record the epidural epileptiform activity induced by chemical administration. The chemically inert nature of FPs and the gold electrode is expected to

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facilitate reliable neural interfacing without abiotic issues. Plasma-treatment-induced covalent binding of FP films can also be utilized in a variety of applications requiring durability, such as implantable biosensors and sensor platforms operating under chemically harsh environments including humid conditions.

INTRODUCTION The failure of the interfaces between implanted neural electrode arrays and neural tissues remains a major obstacle to achieving chronic control of prosthetic limbs and external devices via the long-term stable extraction of neural information from the central and peripheral nervous system.1-3 Interface failures are classified as biotic and abiotic mode failures, whereby devicerelated or system-related failures, respectively, result from biological or tissue responses and disruptions of the blood-brain barrier that may reduce the ability to record neural activity.4 The chronic responses of neural tissues to implanted electrode arrays have been considered the leading cause of neural signal degeneration over time, and many factors such as the electrode size, shape, material stiffness, surface roughness, porosity, and chemical modification contribute to the response.5 For example, substantial mechanical mismatch between the nervous tissue and neural electrodes has been hypothesized and demonstrated to activate foreign body reactions.6,7 Recently, Barrese et al.8,9 investigated in detail the failure of Si-based intracortical microelectrode arrays and reported that electrode corrosion and cracking and delamination of the passivation layer caused long-term reductions in impedance and signal quality. They also suggested that oxygen radicals released by cells of the immune system may play a major role in electrode degradation.


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Polymeric neural prosthetic devices have attracted much attention because they do not break easily under mechanical stress and can accommodate organ curvature. Reduced mechanical mismatch between polymeric devices and brain tissues, attributed to the relatively low Young’s modulus of these devices compared to that of conventional metals and Si, may alleviate stress on the biotic issues.10,11 Currently, several polymers are used in the fabrication of flexible neural prosthetic devices: silicone, PDMS, polyurethane, polyimide, SU-8, parylene C, liquid crystal polymers, and benzocyclobutene.12 Using these polymers, many flexible neural prosthetic devices have been developed for in vivo interfacing with the cortex, retina, spinal cord, and nerves.13-17 However, although the durability of polymer-based neural electrodes has been reported,18, 19 successful long-term reliable interfacing has not yet been achieved. Stieglitz has commented on the longevity and reliability of polymeric neural prosthetic devices and noted the importance of the electrode-polymer interface.20 In almost all of the developed flexible devices, a Cr or Ti layer has been introduced to promote adhesion for typical electrode metals, such as Au and Pt, because of their inherently weak interactions with polymers. There are abundant Na+ and K+ ions in the nervous system and cerebrospinal fluid, which can diffuse into nanocracks within the metal layer and into the gap between the metal and the passivation layer. The diffused ions can dissolve the adhesion promoter or weaken the adhesion between the adhesion promoter and the substrate and/or passivation layer, ultimately resulting in the detachment of the electrode from the substrate. This problem becomes more prominent when the water permeability of the passivated substrate is high. In this regard, fluoropolymers (FPs) are promising candidates for the fabrication of reliable neural electrodes, owing to their hydrophobicity, very low water absorption, chemical inertness against acids and bases, and biocompatibility.21


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In general, flexible prosthetic devices are fabricated by a metallization process and a passivation process. The metallization of FPs has long been reported, and enhanced adhesion between FPs and an evaporated Au layer was achieved without adhesion metals using FP plasma treatment. The primary effect of plasma treatment has been interpreted as an improved mechanical interlocking due to increased surface roughness.22-24 In 1996, Hyland et al.24 attempted to fabricate a subdural flexible electrode array by patterning a Au film e-beamevaporated onto radio frequency (RF) plasma-treated polytetrafluoroethylene (PTFE). They then deposited 200-nm-thick SiO2 onto the Au-patterned PTFE film by e-beam evaporation for passivation. Although they opened the active electrode surface and modified the electrode to have lower impedance, they did not report improved performance or durability. With FP materials, the evaporated SiO2 would be unsuitable for passivation because of its rigidity, high number of pinholes or cracks, and poor adhesion with FPs. In this work, we performed passivation based on FP characteristics, and fabricated flexible neural prosthetic devices solely composed of FP and Au, which are both chemically inert. The devices contain no degradable components and would therefore eliminate at least abiotic issues and thus provide a more reliable neural interface. In this work, the direct adhesion of plasma-treated FP films was investigated at temperatures below the melting point without using pastes or ligands. Although conventional melt-driven processes can be used to adhere FP films, these processes are inevitably accompanied by lateral and vertical deformation of the metal patterns, which is increasingly detrimental as the size of the electrode decreases. Furthermore, in the absence of oxygen, noble gas plasmas, such as argon or helium, are well-known to break C-C, C-H, and C-F bonds and generate free radicals at the polymer surface.25 The generated free radicals can react with each other, leading to branching or


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crosslinking. Uncharged Ar particles and vacuum ultraviolet (UV) photons have been shown to play a key role in crosslinking;26 moreover, plasma-generated free radicals have been revealed to survive for several days and have been identified as the underlying source of plasma-grafting copolymerization-based surface modification of FPs for biomedical applications.27 However, to the best of our knowledge, there have been no reports on the co-adhesion of plasma-treated FP films for biomedical applications. Here, we report the fabrication of a neural prosthetic electrode that does not require an adhesion-promoting metal layer, in which the FP film is both a substrate and passivation material and Au is the electrode material.

EXPERIMENTAL METHODS RF Plasma Treatment of FP Films and Thermal Pressing. The fluorinated ethylene propylene (FEP) film was selected for its relatively low glass transition temperature and melting point and for its optical transparency. FEP has been used to encapsulate metal lead wires for intramuscular electrodes28 and as a substrate for flexible micro-antennae for in vivo magnetic resonance imaging.29 RF magnetron plasma treatment and thermal pressing were employed to generate radicals and to crosslink the generated radicals, respectively. All FEP films used in this study were purchased from CS Hyde (USA). Prior to plasma treatment, FEP films were cleaned with ethanol under sonication followed by drying with N gas. A set of two FEP films was treated with Ar RF plasma and then placed between pressing plates with the two plasma-treated surfaces facing each other. After thermal pressing, the adhesion strength was determined from the 180° peel test. For the co-adhesion study, 5 cm × 5 cm × 75 µm (A) and 1 cm × 15 cm × 50 µm (B) FEP films were treated with Ar RF plasma. A commercial 8” magnetron gun (Angstrom Sciences, USA) was used for the Ar RF plasma treatment. The samples were then placed


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between two pressing plates with the Ar RF plasma-treated surfaces facing each other for 3 min, thermally pressed for 20 min at the desired pressing temperature and pressure, and subsequently cooled to a temperature below the glass transition temperature of FEP. After thermal pressing, FEP film B was peeled from FEP film A. A 180° peel test was performed using a commercially available peel tester BMSPT-180PT (BMS Tech, Korea). The pulling speed during the peel test was 10 mm min-1. Surface Characterization of RF Magnetron Plasma-treated FEP. The RF plasma-treated FEP surface was characterized based on static water contact angle (WCA) measurements, atomic force microscopy (AFM), and field emission scanning electron microscopy (FESEM). The WCA was measured using a CAM100 (KSV Instruments, Finland) immediately after plasma treatment under ambient conditions. Measurements were performed with deionized 0.5 µL water droplets at eight different positions, and the WCA measurement limit was approximately 160°. AFM and FESEM images were obtained using an NX10 AFM (Park Systems, Korea) and an SU5000 FESEM (Hitachi, Japan), respectively. Fabrication of FEP-based Neural Prosthetic Electrode. Scheme 1 shows the unit fabrication processes for the proposed FEP-based neural prosthetic devices. (1) A 127-µm-thick and 50 mm × 50 mm FEP film was sandwiched between two 2.5-mm-thick quartz plates and then placed between press plates. The quartz plates were used to remove scratches. The FEP film was then thermally pressed at temperatures of 250–265 °C and at a pressure of 2 bar cm-2 for 30 min. (2) The pretreated FEP films were then exposed to Ar RF plasma at a working pressure of 15 mTorr and RF power of 40 W for 4 min. (3) A 50–60-nm-thick Cr layer was thermally deposited on the 600–1,000-nm-thick Au layer and used as a sacrificial layer, not as an adhesion layer. Wet Au etchant (Gold Etchant TFA, Transene, USA) and Cr etchant (CR-7, Cyantek, USA) were used to


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pattern Au and Cr, respectively. (4) The Au-patterned FEP film and a 12.7-µm-thick FEP passivizing film were placed side-by-side and treated by Ar RF plasma for 30 s under RF power of 90 W and working pressure of 15 mTorr. (5) The two samples were thermally pressed with the Ar RF plasma-treated surfaces facing each other for 20 min, under a pressing temperature of 200 °C, and pressure of 8 bar cm-2. (6) A 300-nm-thick Al layer was thermally deposited and wet Al etchant (AL-12S, Cyantek, USA) was used to pattern Al. (7) Oxygen RF magnetron plasma etching was applied for approximately 1 h under RF power of 75 W and working pressure of 15 mTorr. (8) A porous Au structure was formed and IrOx was electrochemically deposited. To produce the nanoporous Au structure, an electro-co-deposited Ag-Au electrode was immersed in concentrated nitric acid and maintained at 70 °C for 1 h. Details about the electrochemical codeposition of the Ag-Au alloy and electrochemical deposition of IrOx have been previously published.30

Scheme 1. Unit fabrication processes for FEP-based neural prosthetic devices.


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Rat Epidural Implantation. Seven Sprague-Dawley rats (250–300 g; Samtako Co., Korea), which arrived 1 week before the experiment, were kept in a breeding environment maintained at 21 ± 2ºC and 40–60% humidity under a 12-hour light-dark cycle. All experimental procedures were conducted in strict accordance with procedures approved by the Institutional Ethical Committee on Experimental Use of Animals at Wonkwang University. Rats were anesthetized with an intraperitoneal injection of urethane (10 mL kg-1 of 20% urethane dissolved in normal saline). The head was then fixed in a stereotaxic frame under spontaneous respiration, and 2% lidocaine was injected into the scalp to prevent possible sensations of pain. After shaving the top of the head, a skin incision was made on the skull in the rostro-caudal direction. The soft tissue on the skull was removed and the bregma (reference point) was identified. Under a surgical microscope, a craniotomy was performed by drilling through the right side parietal bone region with a dental drill, and then the dura mater covering the cortex was exposed. The 16-channel electrode array was tightly positioned on the exposed dura mater using a micromanipulator (SM21, Narishige, Japan). A stainless needle electrode as a reference was positioned on the nasion area. Throughout all experiments, the body temperature was maintained at 37.5 °C using a closed-loop animal blanket system (FHC, USA), and the response to pain and spontaneous whisking was monitored. Additional urethane was applied if rats developed spontaneous whisking or showed response to pain. In vivo Recording of Electrocortical Signals. The electrocortical signals from the epidural electrocorticography (ECoG) were amplified using a custom-made headstage amplifier and a second amplifier (×10,000) (CyberAmp 380, Axon Instruments, USA). Because of the limited number of channels in the headstage amplifier, the signals were recorded from only three recording electrodes out of the 16-channel electrode array. The signals were filtered from 0.5 Hz


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to 100 Hz by implementing a digital filter in the CyberAmp. The amplified and filtered signal was transmitted to an oscilloscope (Tektronix 5113, USA). The signals were saved to a personal computer using an analog-to-digital converter (CED 1401, Cambridge Electronic Design, UK) and the software Spike 2 (version 7.0, Cambridge Electronic Design, UK). The baseline electrocortical signals were recorded for another 1 h, and then paroxysmal ECoG discharges were induced by direct local application of the chemoconvulsant picrotoxin into the exposed dura mater of the brain using a micro-syringe with a volume of 100 µL. Picrotoxin was dissolved into artificial cerebrospinal fluid with a concentration of 1 % by volume.

RESULTS AND DISCUSSION

Figure 1. (a) Typical strain-load curves obtained from Ar RF plasma-treated FEP films (blue and violet lines) and untreated FEP films (green line). (b) Adhesion strength (solid circles) and WCA (open circles) with respect to the hydrogen feed ratio during Ar RF plasma treatment (RF power: 70 W; working pressure: 15 mTorr; pressing temperature: 200 °C; and pressing pressure: 8 bar cm-2).


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Figure 1a shows typical examples of strain-load curves obtained for FEP films treated with Ar RF plasma and untreated FEP films. As expected, almost no adhesion was detected from the thermally pressed untreated FEP films and when only one FEP film was plasma-treated. Adhesion occurred only when both FEP films were treated with Ar RF plasma, indicating that the adhesion is induced by the thermal crosslinking of radicals generated by the Ar RF plasma treatment. For reference, the adhesion strength of Scotch® tape (810D) and cleaned glass was measured to be approximately 0.9 N/cm. To confirm the validity of thermal crosslinking as an underlying mechanism for the adhesion of Ar RF plasma-treated FEP films, hydrogen gas was introduced as a radical scavenger. The hydrogen was expected to decompose into hydrogen radicals and then recombine with surface radicals generated by Ar RF plasma, leading to a stable end-capping of broken chains and hence reduced crosslinking capability. Figure 1b shows the effect of the hydrogen gas feed ratio on the adhesion strength (solid circles) and WCA (open circles). As shown in the figure, the adhesion strength decreased with an increasing hydrogen feed ratio, as expected. The WCA also decreased with an increasing hydrogen gas mixing ratio, implying that hydrogen atoms were substituted into the FEP surface. The substituted hydrogen atoms would be expected to contribute to an increased surface free energy, resulting in a lower WCA. These results support the suggestion that the underlying mechanism for the co-adhesion of Ar RF plasmatreated FEP films is the thermal crosslinking of radicals. Radicals generated by hydrogen elimination have previously been identified as the major source of radicals, according to an electron spin resonance study.31 Because the properties of the plasma-treated surface are influenced by the plasma conditions, to determine the optimal RF plasma treatment conditions, we briefly investigated the


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effect of plasma control parameters such as the RF power, working pressure, and treatment time in terms of the adhesion strength, WCA, and surface morphology. In addition, we observed the effect of the pressing temperature and pressure on the adhesion strength. As shown in Figure 2a, the adhesion strength and WCA increased steeply with increasing Ar RF plasma treatment time at the incipient stage for up to approximately 1 min, indicating that short-term treatment below 3 min is sufficient to promote the co-adhesion of FEP films. After the steep increase, both the adhesion strength and WCA become saturated. When the treatment time exceeded 7 min, the

Figure 2. (a) Effect of RF plasma treatment time on adhesion strength (solid circles) and WCA (open circles) under the following conditions: RF power: 40 W; working pressure: 15 mTorr; pressing temperature: 200 °C; and pressing pressure: 8 bar cm-2. (b) FESEM and AFM images of the FEP film treated by Ar RF plasma for 30 s, 4 min, and 8 min. AFM imaging was performed to check for possible damage during FESEM imaging. All the FESEM and AFM images are shown in Figure S1.

FEP surface became super-hydrophobic (>160°), although the adhesion strength remained almost the same. The intact contact angle of 114° was very close to that reported by Ferchichi et al.32 for


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FEP. The treatment time dependence of the adhesion strength of Ar RF plasma-treated FEP films was quite similar to that of O and He plasma-treated FEP films.33 The initial steep rise was attributed to the formation of long-lived trapped radicals generated by the plasma treatment. The saturation of the adhesion strength may indicate that the radicals were generated within a shallow region very close to the surface. The FESEM and AFM images shown in Figure 2b reveal that the FEP surface changed from a lamella-like structure to a so-called protrusion structure34 as the treatment time increased. These results indicate that the adhesion strength of the Ar RF plasmatreated FEP is not sensitive to the surface morphology. Considering the probability of mechanical interlocking, the roughened morphology may favor the adhesion of Au and FEP more than the lamella-like structure, unless excessively thick Au layers are required to fill the roughened surface. The super-hydrophobic protrusion structure would be useful for transparent self-cleaning, anti-fouling, and anti-sticking surfaces.35, 36

Figure 3. (a) Effect of RF plasma treatment time on adhesion strength (solid circles) and WCA (open circles) under the following conditions: working pressure: 15 mTorr; treatment time: 30 s; pressing temperature: 200 °C; and pressing pressure: 8 bar cm-2. (b) FESEM and AFM images of


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the FEP films Ar RF plasma-treated at RF powers of 20, 90, and 160 W. All the FESEM and AFM images are shown in Figure S2.

As shown in Figure 3a, the adhesion strength (solid circles) increased with increasing RF power at lower RF powers and decreased after exhibiting a maximum adhesion strength at approximately 90 W. As previously reported, when the RF power is low, the generation of radicals dominates over chemical etching and results in an increase in the adhesion strength; as the RF power increases further, chemical etching becomes dominant over radical generation and leads to a decrease in the contacting interface area, ultimately resulting in a decrease in the adhesion strength.25 In support of this interpretation, the FESEM and AFM images shown in Figure 3b reveal the transition from a lamella-like structure to the protrusion structure32 with increasing RF power.

Figure 4. (a) Effect of working pressure on adhesion strength (solid circles) and WCA (open circles) under the following conditions: RF power: 70 W; treatment time: 30 s; pressing temperature: 200 °C; and pressing pressure: 8 bar cm-2. (b) FESEM and AFM images of the FEP films Ar RF plasma-treated under working pressures of 15, 35, and 50 mTorr.


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Figure 4a shows that the adhesion strength (solid circles) and WCA (open circles) are largely independent of the working pressure below 30 mTorr, and then they decrease between 30 and 35 mTorr, and then become relatively constant again above 35 mTorr. This result indicates that a lower working pressure favors higher adhesion strength. The reduction of the adhesion strength and WCA may have been caused by the reduced mean free path of the plasma at higher working pressures and consequent reduced kinetic energy of the plasma gas. As shown in Figure 4b, a lamella-like morphology can be seen clearly in the FESEM and AFM images of the FEP film treated at 15 mTorr, but only faintly for those treated at 35 and 50 mTorr. This result confirms the reduced kinetic energy of the plasma gas at a higher working pressure. In conjunction with the RF power dependence of the adhesion strength, this result indicates that the ion-beam irradiation or ion-beam-assisted RF plasma technique could be expected to generate more radicals from deeper regions owing to their inherent higher kinetic energies, leading to enhanced adhesion strength.


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Figure 5. Adhesion strength with respect to the pressing temperature (RF power: 90 W; working pressure: 15 mTorr; treatment time: 30 s; and pressing pressure: 8 bar cm-2).

Figure 5 shows the pressing temperature dependence of the adhesion strength of Ar RF plasma-treated FEP films. The unfilled circles represent results obtained from untreated FEP films. The adhesion strength of FEP increased linearly with increasing pressing temperature and exceeded that of Scotch® tape (810D) and glass (approximately 0.9 N/cm) at 180 °C, which is appreciably below the melting point of FEP. The FEP adhesion strength abruptly increased above 240 °C, including the adhesion strength of the untreated FEP films, presumably because of the thermally-induced interlocking of chains. Above 250 °C, the FEP films were molten and did not peel. We also investigated the effect of pressing pressure on the adhesion strength. As shown in Figure S3, the adhesion strength increased at lower pressing pressures and was insensitive to pressing pressures higher than 5 bar cm-2.


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Figure 6. Fabrication of FEP-based epidural ECoG electrode array: (a) Design and layout of the epidural ECoG electrode array. (b) FESEM image of the thermally-deposited Au layer surface. (c) FESEM images of the oxygen RF magnetron plasma-etched FEP-based ECoG electrode array and the magnified etch pattern with a high aspect ratio. Detailed etching conditions are presented in Figure S4. (d) Image of the fabricated epidural ECoG electrode array and the fabricated array connected to a 16-pin header male connector via a 16-pin flexible flat cable connector. (e) Cyclic voltammetry (CV) results for the untreated Au electrode. (f) FESEM images of the IrOx-


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electrodeposited nanoporous Au electrode. (g) Electrochemical impedance spectroscopy results for the untreated Au, nanoporous Au, and IrOx-electrodeposited nanoporous Au electrode. (h) Comparison of the CV results from untreated Au, nanoporous Au, and IrOx-electrodeposited nanoporous Au electrodes. (i) Voltage transient of the IrOx-electrodeposited nanoporous Au electrode recorded with increasing biphasic pulse amplitude (100 µs duration and 30 µs interpulse delay).

Subsequently, we fabricated an in vivo 16-channel epidural ECoG electrode array based on plasma treatment-induced enhancement of FEP and Au and covalent binding of FEP films. Figure 6a shows the layout of the epidural ECoG electrode array. The diameter of the active electrode was 100 µm. First, the FEP film was thermally pressed to relax the stresses and remove scratches formed during the film manufacturing process. When this pretreatment step is omitted, scratches often disconnect the inter-connection lines and stress inevitably induces film shrinkage, which in turn causes pattern misalignment. The pretreated FEP film was then exposed to Ar RF plasma, followed by sequential thermal deposition of a 1-µm-thick Au layer and 50–60-nm-thick Cr layer. Figure 6b shows a typical FESEM image of the thermally-deposited Au layer. The adhesion of the Au layer and plasma-treated FEP was optically examined using Scotch® tape (810D), and the Au layer was not observed to become detached over the course of this work. The active electrodes, contact pads, and connection lines were then patterned by employing conventional lithography and wet etching of Cr and Au. In this step, the Cr layer was patterned to partially cover the active Au electrode and Au contact pads. Both the Au-patterned FEP surface and a 12.7-µm-thick FEP passivation film were then exposed to Ar RF plasma. They were then thermally pressed at 200 °C with the Ar RF plasma-treated surfaces facing each other.


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Next, to access the active electrodes and connection pads, an aluminum layer was deposited by thermal evaporation and patterned by conventional lithography, and the sample was then exposed to oxygen RF plasma. Although the dry etching of FEP using oxygen RF magnetron plasma resulted in an etch pattern with a high aspect ratio (Figure 6c), the technique also resulted in the formation of metal fluoride (MFx). For example, the fluorine radicals generated during the dry etching were prone to form gold fluoride (AuFx), which is highly stable and was not soluble in many chemicals and could even halt the etching process.37 AuFx is electrically insulating and can thus result in quite low current densities. Delivering a higher RF power (approximately 400 W) to the Ar plasma may remove the AuFx layer;38 however, this would also etch the gold layer and may result in thermal damage. To avoid this problem, we introduced Cr as a sacrificial layer both on the Au electrodes and on the contact pads. Although CrFx was also formed, it was easily removed by dissolving intact Cr using an etchant with sonication. Figure 6d shows the fabricated epidural ECoG electrode array connected by a 16-pin flexible flat cable connector to an interface board for electrochemical deposition and characterization. Figure 6e shows a typical kineticscontrolled cyclic voltammetry (CV) curve, with the characteristic anodic and cathodic peaks that were obtained from the intact Au electrode. To minimize noise in the neural signal recording and to ensure efficient charge injection capabilities in the Au electrode, the electrode was modified with an IrOx-electrodeposited nanoporous Au structure. A Ag:Au alloy was first electrodeposited, and then the electrode was immersed in concentrated nitric acid and maintained at 70 °C for 1 h to remove Ag, followed by the electrodeposition of IrOx. The FESEM images shown in Figure 6f confirm the successful surface modification with the nanoporous structure. The detailed conditions for the electro-codeposition of the Ag:Au alloy, the electrodeposition of IrOx, and various electrochemical


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measurements, including electrochemical impedance spectroscopy, CV, and voltage transient tests, are described in the references.30,39 The electrochemical impedance and charge storage capacitance derived from the CV of the IrOx-electrodeposited nanoporous Au electrode were 3.1 ± 0.1 × 10-3 Ω (at 1 kHz, Figure 6g) and 4.7 mC cm-2 (Figure 6h), respectively. The charge injection limit of the modified electrode derived from voltage transient tests (Figure 6i) was 1.2 mC cm-2. The charge injection efficiency, defined as the ratio of the charge injection limit to the charge storage capacitance, was recorded as 25.5%, almost twice that reported for sputtered IrOx films and electrodeposited IrOx films, confirming the role of nanoporous Au in the efficient charge transfer from the Au electrode to IrOx through the large surface area and in the more efficient charge injection. The nanoporous structure may contribute to the enhanced charge injection efficiency by providing increased charge delivery from the interface to the bulk electrolyte. The chemical stability of the Au-FEP interface can be inferred from the chemical Ag leaching process using concentrated nitric acid at 70 °C for 1 h. In previous reports on the modification of a Au multi-electrode array (MEA) with a nanoporous Au structure,30,39 the nanoporous Au was observed to peel off from the ITO when prolonged leaching, e.g., 20 min, was attempted for complete Ag removal. This was probably caused by the dissolution of the Cr or ITO, which were used as an adhesion promoter and as a base electrode for interconnection, respectively. In the present FEP-Au interface system, no adhesion promoter was used, and therefore the complete removal of Ag could be achieved by prolonged leaching. The typical CV curve (red line) of the nanoporous Au electrode shown in Figure 6h demonstrates the chemical stability of the FEP-Au interface system. Considering the chemical inertness of FEP and Au, chemical interactions between the plasma-treated FEP and Au cannot be expected to occur under


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the fabrication conditions employed. The samples were immersed overnight in a Cr and Al etchant with acetone during the fabrication process; however, no mechanical or electrochemical damage was found. Because both FEP and Au are also inert to alkaline ions, no separation of the Au from the FEP film was expected when it was exposed to body fluids, such as cerebrospinal fluid, as long as the temperature of the solution did not exceed the glass transition temperature of FEP. The adhesion of Ar RF plasma-treated FEP and evaporation-deposited Au was largely attributable to physical anchoring owing to the surface roughening induced by the RF plasma treatment and the penetration of hot Au atoms during the thermal evaporation. In terms of durability, the long-term electrochemical corrosion of IrOx is anticipated;40 however, the nanoporous Au electrode can take over the function of IrOx, although the charge injection limit of nanoporous Au is slightly lower than that of IrOx..

Figure 7. Recording of epileptic discharge: (a) Photograph of the FEP-based ECoG MEA positioned on the Sprague-Dawley rat dura meter. (b) Epidural baseline and epileptic discharge signals induced by picrotoxin administration.


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To evaluate the in vivo electrocortical signal recording performance of the developed ECoG MEA, epidural ECoG recording experiments were performed on Sprague-Dawley rats. After the surgical procedure (Figure 7a), baseline electrocortical signals were recorded for 1 h and then paroxysmal ECoG discharges were induced by direct local application of picrotoxin, a chemoconvulsant to the brain. During baseline epidural ECoG recording, an irregular oscillation of the cortical local field potentials was observed to have an amplitude of 330 ± 110 µV. A single direct application of 1% picrotoxin to the exposed dura mater resulted in epileptiform activity approximately 3–5 min after treatment. The drug dosage was determined according to Bures et al.41 As shown in Figure 7b, the activity reached a constant level within 10–15 min of administration. Signs of epileptiform activity persisted for approximately 3 h. The mean spike frequency and amplitude of epileptiform activity were 38.5 ± 12.7 spike/min and 1,150 ± 250 µV, respectively, from channel 14. The three recording electrodes showed very similar patterns of baseline cortical activity and epileptiform activity. These results confirm the epidural ECoG signal recording performance of the developed FP-based electrode array. The combination of the developed FP-based electrode array with any suitable electrode for deep brain stimulation would enable the construction of a closed-loop neuromodulator.42

CONCLUSION We have covalently bound FEP films by thermally pressing Ar RF plasma-treated FEP films at temperatures below the material’s melting point. The adhesion strength achieved at the pressing temperature of 220 °C was twice that of Scotch® tape (810D) and glass. We have also demonstrated the fabrication of a FP-based Au ECoG electrode array composed solely of FEP and Au by applying Ar RF plasma-induced adhesion enhancement to FEP and Au, covalent


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binding of Ar RF plasma-treated FEP films, and oxygen RF plasma etching of FEP. The Au electrode was modified with electrodeposited IrOx on a nanoporous Au structure. The epidural ECoG signal recording performance of the FEP-Au interface-based ECoG electrode array was evaluated in vivo and was verified to record the epileptiform activity induced by chemical administration. The chemical inertness of the fabricated FEP-Au interface-based electrode array was confirmed during a Ag leaching process performed in hot concentrated nitric acid at 70 °C for 1 h. In terms of addressing abiotic failure issues, the FP-Au interface-based neural prosthetic device offers much promise for reliable neural interfacing owing to its lack of degradable components and its chemical inertness. In addition to neural prosthetic devices, RF plasma treatment-induced covalent binding of FP films can be applied to a variety of applications requiring durability, such as implantable biosensors, sensors applied in extreme chemical environments, interconnect platforms for wearable electronics, and all FP-based microfluidic devices.43,44 In addition to RF plasma, a wide range of techniques is available for generating radicals, such as ion-beam and ultraviolet irradiation, electron bombardment, glow discharge, and remote plasmas, and combinations of these techniques with advanced fluorine chemistry could further enhance the co-adhesion of FPs. Very recently, Minev et al.7 reported that a soft neural implant triggered foreign-body reaction and Shen et al.45 demonstrated that extracellular matrix-based implantable neural electrodes matching the elastic modulus of brain tissue can reduce the inflammatory strain field in the tissue. Considering the established surface modification capability of FPs through graft polymerization46-48 and advances in fluorocarbon chemistry, the FP-based neural prosthetic device may have the potential to solve biotic issues as well.


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ASSOCIATED CONTENT Supporting Information Figures showing FESEM and AFM images of Ar RF plasma-treated FEP surfaces and adhesion strengths of Ar RF plasma-treated FEPs with respect to the pressing pressure are included in the supplemental file. AUTHOR INFORMATION Corresponding Author *E-mail: [email protected] Author Contributions All authors contributed to this work and the writing of this manuscript. ACKNOWLEDGMENT This work was supported by ETRI (17ZH1100) and by the Pioneer Research Center Program through the National Research Foundation of Korea, funded by Ministry of Science and ICT (2012-0009464). The authors have no competing financial interests to declare.


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Table of Contents Graphic


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Fluoropolymer-Based Flexible Neural Prosthetic Electrodes for

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