Implantable Graphene-based Neural Electrode Interfaces for

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Implantable Graphene-based Neural Electrode Interfaces for Electrophysiology & Neurochemistry in In Vivo Hyperacute Stroke Model Ta-Chung Liu, Min-Chieh Chuang, Chao-Yi Chu, Wei-Chen Huang, HsinYi Lai, Chao-Ting Wang, Wei-Lin Chu, San-Yuan Chen, and You-Yin Chen ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.5b08327 • Publication Date (Web): 14 Dec 2015 Downloaded from http://pubs.acs.org on December 20, 2015

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ACS Applied Materials & Interfaces

Implantable Graphene-based Neural Electrode Interfaces for Electrophysiology & Neurochemistry in In Vivo Hyperacute Stroke Model Ta-Chung Liu1, Min-Chieh Chuang2, Chao-Yi Chu1, Wei-Chen Huang3, Hsin-Yi Lai4, Chao-Ting Wang5, Wei-Lin Chu6, San-Yuan Chen1,* and You-Yin Chen5,* 1

Department of Materials Science and Engineering, National Chiao Tung University, No. 1001, Ta-Hsueh Rd., Hsinchu, Taiwan 300, ROC 2 Department of Chemistry, Tunghai University, No. 181, Sec. 3, Taichung Port Rd., Taichung ,Taiwan 407, ROC 3 Department of Materials Science and Engineering, Carnegie Mellon University, No.5000 Forbes Avenue, Wean Hall 3325, Pittsburgh, PA 15213, USA 4 Interdisciplinary Institute of Neuroscience and Technology, Zhejiang University, No.866 Yuhangtang Rd., Hangzhou, Zhejiang Province, 310058, China 5 Department of Biomedical Engineering, National Yang Ming University, No.155, Sec. 2, Linong St., Taipei, Taiwan 112, ROC 6 Department of Biomedical Engineering, University of Michigan, No.2200 Bonisteel Boulevard, Ann Arbor, MI 48109-2099, USA *

Correspondence should be addressed to either of the following: Prof. San-Yuan Chen Department of Materials Science and Engineering, National Chiao Tung University, No. 1001, Ta-Hsueh Rd., Hsinchu, Taiwan 300, ROC Email: [email protected]; Prof. You-Yin Chen Department of Biomedical Engineering, National Yang Ming University, No.155, Sec. 2, Linong St., Taipei, Taiwan 112, ROC Email: [email protected]

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Abstract Implantable microelectrode arrays have attracted considerable interest due to their high temporal and spatial resolution recording of neuronal activity in tissues. We herein presented an implantable multichannel neural probe with multiple real-time monitoring of neural-chemical and neural-electrical signals by a non-enzymatic neural-chemical interface, which was designed by creating the newly developed reduced graphene oxide- gold oxide (rGO/Au2O3) nanocomposite electrode. The modified electrode on the neural probe was prepared by a facile one-step cyclic voltammetry (CV) electrochemical method with simultaneous occurrence of gold oxidation and GOs reduction to induce the intimate attachment by electrostatic interaction using chloride ions (Cl-).The rGO/Au2O3-modified electrode at a low deposition scan rate of 10 mVs-1 displayed significantly improved electrocatalytic activity due to large active areas and well-dispersive attached rGO sheets. The in-vitro amperometric response to H2O2 demonstrated a fast response of less than 5s and a very low detection limit of 0.63 µM. In in-vivo hyperacute stroke model, the concentration of H2O2 was measured as 100.48 µM ± 4.52 µM for rGO/Au2O3 electrode within one hour photothrombotic stroke, which was much higher than that (71.92 µM ± 2.52 µM) for non-coated electrode via in-vitro calibration. Simultaneously, the somatosensory-evoked potentials (SSEPs) test provided reliable and precise validation for detecting functional changes of neuronal activities. This newly developed implantable probe with localized rGO/Au2O3 nanocomposite electrode can serve as a rapid and reliable sensing platform for practical H2O2 detection in the brain or for other neural-chemical molecules in vivo.

Keywords: Multichannel neural probe, rGO/Au2O3 nanocomposite, Electrophysiology, Neurochemistry, Stroke

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1.

Introduction Recently, there has been major progress in various implantable biosensors such as

microelectrode arrays for chemical and electrical sensing of the central nervous system (CNS)1. Typically, implantable microelectrode array are used to perform high temporal and spatial resolution monitoring of neuronal activity in tissues2. While the recording of electrical signals can be achieved via the successful realizations of implantable sensor array, proper attention must be paid to integrate electrical signals, with responses acquired from many neural-chemical substances in order to lead to a better identification of neuronal activity. Thus, there is a drive to integrate electrochemical and electrophysiological responses using microelectrode array for recording. Impressively, nanomaterials decorated on microelectrodes give advances in realizing electrical and chemical communications within the nervous system at the micro and nanoscale3, because electrical and chemical neural activity are tightly coupled in the CNS. This indicates that the neural interface between the electrode and neural tissues has played a very important role in the electrical signal intensity of neural implants. The requirements for a successful and implantable microelectrode array must include well intrinsic electrochemical properties4, high electrical performances, and proper biocompatibility because the neural interface was used to sense neuronal chemical and electrical activity (such as neurotransmitters, action potentials, and local field potentials, etc.) through the transduction of electrical charges, which strongly depend on the electrode materials and the surface morphology on the implant probe5-6. So far, many enzyme-based electrochemical/electrical interfaces have been developed in neural detection, because the enzyme-based assays possess good selectivity and sensitivity7-9 but few studies have been focused on neural detection in vivo due to the complex environment in brain. There still exist some practical problems, such as short operational lifetimes and low reusability of these implantable microsensors10. In contrast to enzyme-based assays, inorganic metal oxides and their composites are easily prepared and more stable, even at high temperatures. As a result, non-enzymatic sensors are more suitable for in vivo sensing, because enzymatic sensors are easily 3



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dampened by inflammatory response during implantation. However, one important challenge in developing non-enzymatic sensor in vivo arises from the limitation of low sensitivity and selectivity. These

challenges

directed

us

towards

developing

a

non-enzymatic

and

improved

electrochemical/electrical neural interface, which can be tuned by controlling the architecture of the electrode interface using nanomaterials. Among different nanomaterials, graphene-based material has drawn more attention due to its unique properties and being highly amenable to micro-fabrication11. Moreover, graphene oxide-based (GO-based) materials are electrochemically active toward redox-active species, and achieve intimate connectivity with molecules on their surface due to their excellent electrical conductivity, mechanical properties, chemical stability, and biocompatibility12. More importantly, owing to their solution-processability13, GO-based materials can be readily “functionalized” to enhance catalytic properties14 and reduce thermal noise in recording signals15. However, certain problems such as aggregation, were encountered in assembling a GO-modified metal-based electrode due to the very weak intersheet van der Waals attractions and non-compatible bonding between metal and GO sheets. Many methods, such as the Langmuir-Blodgett deposition16

have

been employed to synthesize single or monolayer GO to resolve the issues and enable the film thickness to be controlled, whereas the disadvantages of weak physical bonding by intermolecular forces and the addition of toxic materials during the manufacturing process must still be overcome. However, these synthesis processes are complicated, because complex chemical reactions are involved. Therefore, in this study, a facile one-step electrochemical method was proposed to fabricate a reduced graphene oxide on the gold electrode by chloride ions. Hydrogen peroxide (H2O2), one of most common representatives of reactive oxygen species in biological system and its accumulation, like other oxidants, has historically been considered a neurotoxic damaging entity that mediates diseases such as trauma and ischemia-reperfusion17-18. Recently, there have been reports to emphasize the importance of graphene-based materials on H2O2 detection19-20. For example, Zhang et al. used graphene quantum dots to sense H2O2 in living cells 4


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with a range of 0.002-8 mM and a detection limit of 0.7 µM21. Furthermore, Maji et al. immobilized gold nanoparticles in mesoporous silica covered graphene oxide sheets to detect H2O2 in cancer cells and found that it could serve as a versatile platform for biosensing.22 In this study, we designed a multichannel neural probe with a non-enzymatic graphene-based electrode to detect the brain’s H2O2 in vivo, while simultaneously integrated with electrophysiology signals for proof-of-concept evaluation of neuronal activity with the photothrombotic stroke model23. The electrode on the neural probe could be modified by the one-step electrophoresis of GO sheets on the gold electrode by chloride ions, which not only acted as an electrolyte in electrophoresis process, but also played an important role as an intermediate cross-linker at the electrode/tissue interface by ion corrosions. Moreover, this kind of deposition method can effectively disperse rGO sheets on the metal electrode to induce the intimate attachment by the electrostatic interaction between etched gold oxide layer and rGO sheets. We believed that this is the first report to demonstrate GO-based deposition on metal electrodes by a facile process combined the cyclic voltammetry (CV) with chloride ion (Cl-) induction as shown in Scheme 1. This process yielded a rough structure with extremely high surface areas and well-dispersed rGO sheets, contributing to a great electrode-electrolyte interface for bio-sensing24. The physiochemical and electrocatalytic properties of the rGO/AuOx electrodes on the neural probe for H2O2 detection were characterized and discussed in a rat brain stroke model in this work. Additionally, somatosensory-evoked potentials (SSEPs) were used to further validate for detecting functional changes of neuronal activities.

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2.

Materials and Methods

2.1. Preparation of rGO/AuOx nanocomposite electrode on the neural probe Electrochemical synthesis was performed by CV using an electrochemical instrument (CHI 614C, CH Instruments, Inc., Austin, TX, USA) with a three-electrode configuration, in which Ag/AgCl served as the reference electrode (No.002243, ALS Co., Ltd, Tokyo, Japan), a platinum wire as the counter electrode (No.002222, ALS Co., Ltd, Tokyo, Japan), and a flexible polyimide-based neural probe (Fabrication process and its respective specifications were shown in Note 1, Supporting Information) with a linear microelectrode array as the working electrode. The neutral aqueous GO suspension (Detailed synthesis process was shown in the Note 2, Supporting Information) was mixed with sodium chloride (NaCl) solution (100 mM) as the electrophoresis medium. The rGO/AuOx nanocomposite was prepared using a scan range of -1.4 V to +1.4 V, within which the gold surface was dissolved at 1.25 V and electro-reduction of GO occurred at -1.0 V25. Additionally, we prepared rGO/AuOx modified electrodes with different deposition scan rates from 10 mVs-1 to 250 mVs-1 for further investigation.

2.2. Characterization of rGO/AuOx nanocomposite electrode on the neural probe The surface morphology and topographic images of the rGO/AuOx nanocomposite electrode on the neural probe was examined using scanning electron microscopy (SEM, JSM S6700, JEOL Ltd., Tokyo, Japan) and atomic force microscopy (AFM). Detailed information can be referred to the Supporting Information Note 3.

2.3. Electrochemical characterizations The electrochemical performance of the rGO/AuOx neural probe was tested using CV in a 0.02 M hydroquinone (HQ)/ 0.1 M phosphate buffered saline (PBS) solution, in which a platinum wire was used as the counter electrode and Ag/AgCl as the reference electrode. The rGO/AuOx neural probe with different deposition scan rates was submitted to analyze by electrochemical impedance 6


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spectroscopy (EIS) during the electrochemical process. The parameters of the equivalent circuits were computed by ZView® software (Scribner Associates Inc. Southern Pines, NC, USA). Finally, the surface kinetics and surface reactions of rGO/AuOx dependent on the deposition scan rates (10, 20, 40, 80, 160, and 320 mVs-1) were examined.

2.4. Interference study and in vitro amperometric experiments As comparison to in vivo, the performance of the rGO/AuOx nanocomposite was validated in vitro using artificial cerebrospinal fluid (aCSF) combined with several electroactive common matters (0.5 mM ascorbic acid (AA), 0.5 mM dopamine (DA), 0.5 mM uric acid (UA), and 1 mM) H2O2 by fast-scan cyclic voltammery (FSCV) at scan rate of 105 Vs-1 to determine optimal reduction potential for H2O2 detection. Detailed information was described in the Supporting Information Note 4.

2.5. In vivo electrophysiological and neurochemical experiments For proof-of-concept evaluation of the in vivo electrophysiological and neurochemical recordings with the rGO/AuOx neural probe, our rGO/AuOx neural probe was used to perform the detection of H2O2 and neural activity in the primary somatosensory cortex (S1FL) of the hyperacute photothrombotic stroke model. Detailed surgical process was described in the Supporting Information Note 5.

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3. Results and discussion 3.1. One-step electrochemical deposition of rGO/AuOx nanocomposite Figure 1 showed the formation steps and the corresponding CV plot of an rGO/AuOx nanocomposite electrode on the neural probe at a range from -1.4 V to 1.4 V at 10 mVs-1. During the first anodic potential scan, a large anodic peak appeared at +1.25 V26. Phase

corresponding

to Reaction Eq. (1) illustrated that the gold electrode was dissolved by chloride ions (Cl-) in a GO/NaCl bath to form a positively charged surface, called the Au3+ surface. Phase corresponding to the reaction Eq. (2) illustrated that during the electrochemical deposition of GO sheets on Au3+ (or Au+) rough surface by electrostatic forces, the oxygen dissociated in water would be anchored in Au3+ (or Au+) to form a hybrid GO/AuOx nanocomposite. Phase

(corresponding

to Reaction Eq. (3), a small cathodic hump appeared at +0.6V, indicating the hydration of gold surface and the formation of Au(H2O)m3+. This provides further evidence for the Au3+ ions formation in the solution during the electrochemical deposition. Phase

presented the

electro-reduction of GO, leading to the formation of a rougher rGO/AuOx nanocomposite. In summary, the negatively charged oxygen functionalities of GO are more attracted towards the positively charged Au3+ ions from the bath solution by electrostatic driving force to promote the intimate attachment, resulting in the formation of rGO/Au2O3 nanocomposite. The following include chemical equations during the one-step electrochemical process: Au + 4Clି → Auଷା + 4Clି + 3eି

Eq. (1)

ା ି 2AuClି ସ + 3Hଶ O → Auଶ Oଷ + 6H + 8Cl

Eq. (2)

ି Au + mHଶ O → Au(Hଶ O)ଷା ୫ + 3e

Eq. (3)

3.2. Comparison of structure characterizations and in vitro electrochemical performance with different scan rate of deposition Figure 2(a) showed the optical image of the neural probe with and without rGO/Au2O3 8


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electrodes and the electrochemical properties and electrical properties of the rGO/Au2O3 nanocomposite electrode are dependent on the surface microstructure, especially for the carbon basal planes with the edge plane defects27 on carbon basal planes, which can be tailored by Cl- ion effect to control the dissolution rate of gold. The CV curves for real-time electrodeposition of rGO/Au2O3 film with different deposition scan rates were shown in the Supporting Information Note 6. The low-magnification and enlarged SEM images of rGO/Au2O3 nanocomposite with different deposition scan rates from 10 mVs-1 to 250 mVs-1 were shown in Figure 2(b)-(d) where a rough microporous rGO-wrapped Au2O3 film structure was observed in the rGO/Au2O3 nanocomposite with a deposition scan rate of 10 mVs-1. Such a microporous microstructure with a few layers of rGO sheets can be attributed to a longer dissolution time and slow oxidation of gold electrode by Cl- ions, causing more small pieces of GO layers to be well-dispersive adhering onto the microporous gold electrode surface. In contrast, at higher deposition scan rates, GO sheets easily aggregated into thick layers on the Au3+ surface, as shown in a high-magnification view in Figure 2(d). Surface analysis by AFM in Figure 2(e)-(f) showed that lots of sharp and smaller size protrusions, produced by chloride ions, were observed in rGO/Au2O3 nanocomposites with 10 mVs-1, while the rGO/Au2O3 nanocomposites with 250 mVs-1 produced flattened and large bulks morphology. Also, the surface area of rGO/Au2O3 nanocomposites with 10 mVs-1, 50 mVs-1 and 250 mVs-1 were estimated as 35.7, 28.5 and 26.5 µm2, and the corresponding average roughness (Ra) were 73.8, 62.0 and 57.7 nm, respectively. The rGO/Au2O3 nanocomposite at different deposition scan rates can be further proven by electrochemical analysis of CV plot, shown in Figure 3(a). The redox current with HQ of the rGO/Au2O3 deposited with 10 mVs-1 was 150 µA, which was higher than the rGO/Au2O3 electrodes with other deposition scan rates, and approximately 3 times higher than the non-coated gold electrode (shown in the Table S3, Supporting Information Note 7). The redox peak current of the rGO/Au2O3 nanocomposite obviously increased with lowering deposition scan rates associated with increased active sites. Furthermore, the peak-to-peak voltage separation (∆Ep = 31 mV) of the 9



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rGO/Au2O3 electrode with 10 mVs-1 was lower than that with other deposition scan rates, and approximately 8 times smaller than that of a non-coated gold electrode. The reason for the rGO/Au2O3 with such a low ∆Ep was attributed to fast electron transfer reactions, resulting from the reduction of GOs into rGOs, and the less aggregated barriers caused by the intimate attachment between rGO and Au2O3. Relative proofs for the surface characterizations of rGO/Au2O3 can be found in Raman spectrum in Figure 3(b) and EIS in Figure 3(c)-(d). Figure 3(b) showed the Raman spectra of rGO/Au2O3 nanocomposite with 10, 50, and 250 mVs-1 deposition scan rates. All spectra exhibited two prominent peaks at 1350 cm-1 and 1583 cm-1, corresponding to the well-documented D and G bands, respectively. The ID/IG ratio usually was used to determine the degree of bond disorder and chain information concerning carbon materials28. As compared to GOs, rGOs exhibited a larger ID/IG ratio, indicative of the restoration of sp2 carbon and a decrease in the average size of the sp2 domains upon the reduction of GO. The ID/IG ratio (ID/IG = 1.245) of the rGO/Au2O3 nanocomposite with 10 mVs-1 was larger than those with 50 and 250 mVs-1 (ID/IG = 1.01 and ID/IG = 0.938), respectively. These results demonstrated that more carboxyl functional groups on the GO plane were partially electro-reduced, because of longer exposure time in reductive currents with decreasing deposition scan rate. As a result, the rGO/Au2O3 nanocomposite formed over lower deposition scan rates retained high electron transfer kinetics due to the reduction of oxides on 2-dimension (2D) carbon plane. EIS results in Figure 3(c) and (d) were illustrated by the Bode plot and the Nyquist plot for rGO/Au2O3 nanocomposite, respectively, at different deposition scan rates of 10, 50, and 250 mVs-1. Inert of Figure 3(d) showed the corresponding mixed kinetic circuit model of rGO/Au2O3 electrode using a constant phase element (CPE) with capacitive behavior modeling the double layer, in parallel with Rct together in series with the electrolyte resistance (Rs). In the Bode plot, the overall resistance was determined by the sum of solution resistance (Rs) and electron transfer resistance (Rct) at low frequencies. Taking Rs as constant for all modified electrodes, rGO/Au2O3 nanocomposite at 10 mVs-1 showed the lowest resistance in Bode plot at low frequencies corresponding to the lowest 10


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electron transfer impedance (Rct, 328 kΩ) found in the Nyquist plot. The values of Rct (the diameter of semicircle) calculated from the Nyquist plot by Zview software in Figure 3(d) were 328 kΩ, 512 kΩ, 578 kΩ and 604 kΩ for the rGO/Au2O3 nanocomposite deposited at 10 mVs-1, 50 mVs-1, 250 mVs-1 and non-coated gold electrode, respectively. The lowest Rct further demonstrated that the rGO/Au2O3 electrode with 10 mVs-1 exhibited the fastest transfer kinetics, owing to well-dispersive deposition manner of rGO sheets. Furthermore, the CPE-P (P defined as phase constant) was calculated to be 0.762, 0.754, 0.726 and 0.578 for rGO/Au2O3 electrode deposited at 10 mVs-1, 50 mVs-1, 250 mVs-1 and non-coated electrode, respectively. The CPE-P value in the range of 0.7 - 0.8 indicated that the rGO/Au2O3 electrode exhibited both resistive and capacitive behaviors, which matched the observations found in Figure 3(a) that faradic behaviors appeared at low applied potentials (-0.05V~0.15V) and capacitive behaviors appeared at high applied potentials (0.15V~0.3V, -0.05V~-0.2V). In addition, according to the CPE-T values (T defined as time constant), double-layer capacitance (Cdl) was estimated as 8.83 × 10-9 F, 3.88 × 10-9 F, 1.26 × 10-9 F and 0.37 × 10-9 F for rGO/Au2O3 electrode at 10 mVs-1, 50 mVs-1, 250 mVs-1 and non-coated electrode, respectively, and listed in the Supporting Information Table S3. The higher Cdl (8.83 nF) implies that heavily charges would be stored within rGO sheets of rGO/Au2O3 electrodes. This proved that the rGO/Au2O3 electrode with 10 mVs-1 possessed large amounts of well-dispersive rGO sheets on the electrode. The electrochemical interface properties of the rGO/Au2O3 nanocomposite electrode with 10 mVs-1 were further verified by CV at different scan rates as shown in the Supporting Information Note 8. The above results demonstrated that the rGO/Au2O3 electrode with 10 mVs-1 displayed best electrochemical/electrical properties fitting as the subsequent in vitro and in vivo amperometric biosensors.

3.3. In vitro H2O2 detection and interference study in aCSF Although electrochemistry appears to be a better tool than microdialysis to detect real-time fluctuations in brain neurochemicals because of its excellent temporal resolution, the specificity of 11



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such

measurements

is

always

problematic.

The

electrochemical

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technique

and

background-substrated fast-scan CV (FSCV) provides chemical selectivity, in addition to high temporal resolution and sensitivity29. With this approach, a cyclic voltammogram is generated to serve as a chemical signature for the analyte of interest, allowing discrimination from other electroactive species in the brain30. In comparison to the in-vivo situation, the effect of common interfering electroactive substances, including AA (0.5 mM), DA (0.5 mM), UA (0.5 mM) and H2O2 (1 mM) in aCSF on FSCV, responses were further assessed and presented in Figure 4(a). Since oxygen is a well-known important interferent during the reduction process of H2O2, the supporting electrolyte usually needs to be deoxygenated by bubbling nitrogen for at least 20 min prior to each experiment31. The reductive peaks were observed around -0.42V, -0.15V, -0.4V, -0.43V and -0.5V for AA, DA, UA, glucose and H2O2 in aCSF, respectively. Here a quite big difference (at least 5-fold) in the reduction peak current was observed between H2O2 and other chemical molecules. The applied voltage of -0.5V in this study distinguished the reduction from those of other electroactive substances commonly found in the brain, and thus the accuracy of H2O2 amperometric responses is less affected. Based on the interference study above, -0.5V was reasonably chosen as the operation voltage to enhance selectivity against the main endogenous brain interference species in an amperometric response. Also, owing to rGO’s ability to react with many phenyl–based molecules by π–π stacking and its excellent electrical properties, rGO-based electrodes are able to simultaneously determine DA, AA, and UA by CV compared to conventional metal-based electrodes32. As a result, the rGO/Au2O3 electrode was well qualified as a non-enzymatic neural-chemical interface due to enhanced selectivity by reducing interference effects in vivo through selecting different applied voltages. In Figure 4(b), an obvious amperometric response (approximately 0.5 nA) appeared when 0.5 mM H2O2 was injected at first at an operation voltage of -0.5 V, and then glucose (1 mM), AA (0.5 mM), UA (0.5 mM) and DA (0.5 mM) injections as interferences produced noises while did not cause observable amperometric changes. When H2O2 (0.5 mM) was added for the second time, the current changed proportionally 12


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(approximately 0.5 nA) even with the existence of the interferents. This demonstrated that the developed rGO/Au2O3 nanocomposite electrode showed a superior selectivity to H2O2, by applying specific voltages.

3.4. In vitro H2O2 calibration Since the surface morphology of GO-Au2O3 nanocmposite electrode was strongly influenced by the deposition scan rate as shown in Figure 5, which will affect the sensing performance of rGO-Au2O3 nanocmposite electrode.

The amperometric response of the rGO/Au2O3 electrode

deposited with 10 mVs-1 performed a very clear stepped shape and a high sensitivity of 449.28 nA/µM･cm2 (r2 = 0.984). In comparison to the rGO/Au2O3 electrode of 50 mVs-1 and 250 mVs-1, the sensitivity was 254.99 and 157.86 nA/µM･cm2 (r2 = 0.981 and r2 = 0.963), respectively. In addition, rGO/Au2O3 electrode of 10 mVs-1 exhibited shorter response time than other two electrodes when larger H2O2 concentration was added because the rGO/Au2O3 nanocomposite deposited with 10 mVs-1 not only induced the direct electron transfer (DET) on the H2O2 electrode but also provided high active sites for H2O2 detection due to the high surface areas of rGO/Au2O3. In contrast, the rGO/Au2O3 nanocomposite deposited with 250 mVs-1 showed the worse sensitivity to H2O2 because of the low surface area, which can be ascribed to the higher absorption of H2O2. Figure 6 showed the amperometric current responses for the rGO/Au2O3 electrodes on the neural probe, where the electrode potential was performed at −0.5 V with N2 saturated in aCSF. The rGO/Au2O3 electrode fabricated with a deposition scan rate of 10 mVs-1 performed a fast response time of 5 sec less than a non-coated gold electrode to reach 100% signal with a detection limit of 0.63 µM (n = 10). The current response of H2O2 in Figure 6(a) presented a very steady-state level within 5 sec after injecting different concentrations of H2O2 from 1 µM to 8 mM. The calibration plot was further shown in Figure 6(b), indicative of the sensitivity of the rGO/Au2O3 electrode at 461.42 nA/µM･cm2 (r2 = 0.984), which is higher than the sensitivity of 221.92 nA/µM･cm2 (r2 =

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0.998) for the non-coated electrode. Such a low detection limit and high sensitivity of the rGO/Au2O3 electrode was closely correlated with large amount of active sites, because of high surface roughness of the modified electrode and the inter-attachment between rGO and Au2O3. Furthermore, the amperometric responses remained quite linear even for high levels of H2O2 in the detection range from 1 mM to 59.31 mM (shown in the Supporting Information Note 9). The rGO/Au2O3 nanocomposite could effectively promote a two-electron coupled with two-proton redox reaction between the H2O2 and the electrode, due to its extraordinary electron transfer and high surface areas.

3.5. In vivo neurochemical and electrophysiological measurements in the hyperacute ischemic stroke To demonstrate that the electrode can distinguish H2O2 when implanted into tissue, background subtracted amperometric response of the oxidation of H2O2 were recorded in the S1FL of a stroke brain. The rGO/Au2O3 modified neural probe was inserted at least 900 µm into a section of brain tissue that encompassed the S1FL brain region. The S1FL region was photothrombotic by laser light of 561 nm. This caused a short interval of cerebral ischemia in rats and induced a highly reproducible damage to the neurons. During the ischemic cascade, lack of oxygen initially caused the neurons to fail to make ATP for energy, and then induced a series of reactions including: 1) calcium ions flowed into the cells to cause the release of lactic acid and glutamate, followed by damaging the cell membrane, 2) harmful biomolecules such as reactive oxygen species (ROS) were produced, and 3) at the same time, a byproduct like H2O2 increased with time during the ischemic cascade. The circumstance of neural system in the brain during ischemic stroke can be real-time diagnosed by integrating significant changes of SSEPs and electrochemical sensing in predicting pre-operative stroke and post-operative stroke by multi-channel implantable neural probe. To further demonstrate it, TTC staining was used to confirm that S1FL brain was hit with ischemic 14


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stroke, where TTC staining resulted in viable tissue being stained as “brick-red”. This coloration was due to a reaction between dehydrogenases in the cells and the infarcted tissue stained as “pale-white”, since they lacked the enzymes with which the TTC reacted. Figure 7(a) showed the stained “pale-white” area corresponded to the S1FL, according to the rat brain atlas33. We also labeled the implanted site (black line) and performed real-time detection of relative changes in H2O2 recorded by the rGO/Au2O3 modified electrode (Channel-3) and a non-coated gold electrode (Channel-2) at -0.5V. Meanwhile, SSEPs were monitored by another electrode (Channel-1) to compare the changes in the neuronal activities before and after stroke. For electrochemical detection as shown in top portion of Figure 7(b), the basal currents were located at nearly scale (-0.42 nA ~ -0.48 nA) for rGO/Au2O3 modified electrode and non-coated gold electrode, because the neural probe was exposed to the same physical and chemical environment and identical applied voltage of -0.5V. The S1FL brain was illuminated by laser light starting at 1500 sec to induce the intravascular obstruction, and then a thrombus was gradually formed in the illuminated blood vessels. The recorded currents reduced with time for both the rGO/Au2O3 modified electrode and non-coated gold electrode. It is reasonable to assume that these differences were attributed to the contribution of produced H2O2 in hypoxia brain. Note that a small reduction of electrochemical current was recorded by the rGO/Au2O3 electrode starting at 1500 sec, while the obvious reduction current was observed after 1800 sec by the non-coated gold electrode due to the poor electrocatalytic to H2O2. For the rGO/Au2O3 electrode, a very sharp reduction of amperometric current of H2O2 (-0.504 nA to -0.613 nA) ranging from 1700 sec to 1800 sec was probable due to a rapid hypoxia caused by the complete formation of thrombus. In our study, the relative amperometric currents (-0.482 nA to -0.796 nA) determined by in vitro calibration ranged from a very low concentration of H2O2 to 100.48 µM ± 4.52 µM (mean ± S.D., n = 3) within one hour. However, for non-coated gold electrode, the relative amperometric currents (-0.448 nA to -0.677 nA) within one hour photothrombotic stroke was estimated up to 71.92 µM ± 2.52 µM (mean ± S.D., n = 3). It is reasonable that the results recorded by rGO/Au2O3 electrode are more reliable, because the 15



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concentration of H2O2 in the cortical or subcortical area was around 100 µM - 150 µM by microdialysis and higher accumulated concentration of H2O2 was found in reperfusion34. For further determining neural integrity, our multi-channel neural probe carried out recording SSEPs simultaneously with H2O2 electrochemical sensing. The waveforms of the averaged SSEPs over the recording sessions were also shown in the bottom of Figure 7(b). Two main SSEP components (recorded from the first bare gold electrode, Channel-1) labeled at the waveform were consistently observed: 1) P1, a peak-patency surface-positive component and 2) N1, a longer latency surface negativity. Significant differences in SSEP amplitudes were observed before and after photothrombotic stroke under the same intensity of forepaw stimuli (P < 0.01, paired t-test, n = 3). The amplitudes of P1 and N1 were 3.14 ± 0.15 mV and 5.53 ± 0.49 mV (n = 3), respectively, during the 200-ms post-stimulus period before the photothrombotic stroke. After induction of photothrombotic stroke, the amplitudes of P1 and N1 appeared at 1.30 ± 0.16 mV and 3.33 ± 0.45 mV during the 200-ms post-stimulus period, respectively. Furthermore, The P1 and N1 latency of SSEP for post-photothrombotic stroke (P1: 54 ± 9 ms and N1: 65 ± 5 ms, n = 3) significantly increased to 69% and 36% higher than the values of pre-photothrombotic stroke (P1: 32 ± 3 ms and N1: 48 ± 5 ms, n = 3), respectively. Compared to the SSEPs changes before photothrombotic stroke, a decrease of neuron discharge reactions and time-delay for nerve conduction were observed after stroke35. In summary, our rGO/Au2O3 neural probe combined with chemical and electrical interfaces provide high sensitivity in H2O2 detection, and furthermore, can for the first time be integrated with the measurement of SSEP changes in S1FL regions pre- and post-photothrombotic ischemia stroke.

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4.

Conclusions We have developed a newly-designed rGO/Au2O3 multichannel neural probe using a facile

one-step electrochemical method through an ion (Cl-)-induced process. The design includes a neural-chemical interface with close electrostatic interaction between the negatively charged GO sheets and positively charged Au+3. The rGO/Au2O3 nanocomposite can provide great fast electron transferring on tissue/electrode interfaces by tuning different deposition scan rates. The high H2O2 sensitivity and low detection limit was found in this rGO/Au2O3 nanocomposite electrode in vitro. Furthermore, our multichannel neural probe serves as a non-enzymatic neural-chemical interface to perform efficient real-time detection of brain H2O2 with electrochemical methods. Moreover, brain dynamics of ischemic strokes can also be simultaneously analyzed with electrophysiology signals measured by SSEPs changes. This implantable device can be used as a rapid and reliable model for studying the brain dynamics by fitting various bio-signals.

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Scheme 1. The in-vivo amperometric experimental setup of laser-inducing ischemic stroke and neural probe with reduced graphene oxide-wrapped gold oxide nanocomposite (rGO/Au2O3) electrode by chloride ion (Cl-)-induced effect.

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Figure 1. Schematic illustration of rGO/Au2O3 modified gold electrode using a one-step electrochemical process. Real-time CV plot for the formation of the rGO/Au2O3 nanocomposite was indicated with the numbers corresponding to the schematic.

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Figure 2. (a) Optical image of the neural probe with and without rGO/Au2O3 electrodes coated. SEM images of the rGO/Au2O3 electrodes with different deposition scan rates during the electrophoresis process: (b) 10 mVs-1, (c) 50 mVs-1, and (d) 250 mVs-1. All rGO/Au2O3 electrodes showed that wrinkled rGO sheets were wrapped onto microporous gold oxide electrodes as shown in the marked enlarged area. (e)-(g) Topographical views of rGO/Au2O3 electrodes at 10 mVs-1, 50 mVs-1, and 250 mVs-1 obtained by atomic force microscopy (AFM) analysis (projected geometric area is 25 µm2). The corresponding measured image surface area and roughness were estimated as 35.7, 28.5, and 26.5 µm2 and 73.8, 62.0, and 57.7 nm, respectively. 20


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Figure 3. (a) CVs of the rGO/Au2O3 electrode with different deposition scan rate of 10, 50, 250 mV-1 and non-coated gold electrode at 0.02 M HQ in PBS at a scan rate of 100 mVs-1. (b) Raman spectra of rGO/Au2O3 nanocomposite formed over different deposition scan rates. All spectra of rGO/Au2O3 electrodes showed the D-band (at 1360 cm-1) and the G-band (at 1680 cm-1). (c) Bode plot of EIS results for rGO/Au2O3 deposited at 10 mVs-1, 50 mVs-1, 250 mVs-1, and non-coated gold electrode. (d) Impedance loci (Nyquist plots) of rGO/Au2O3 deposited at 10 mVs-1, 50 mVs-1, 250 mVs-1, and non-coated gold electrode.

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(b)

(a) 10 8 6 4 2 0 aCSF -2 AA -4 DA -6 UA H2O2 -8 -10 -1.5 -1.2 -0.9 -0.6 -0.3 0.0 0.3 0.6 0.9

0.0 Current(nA)

Current(nA)

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AA 0.5mM

Glucose 1mM H O 0.5mM 22O22 0.1mM

-0.5

-1.0

DA 0.5mM

UA 0.5mM

HH2O 0.1mM 2O22 0.5mM

-1.5 0

100 200 300 400 500 600 700 800

E/V

Time(sec)

Figure 4. (a) FSCVs recorded with rGO/Au2O3 electrode at AA (0.5 mM), DA (0.5 mM), UA (0.5 mM), and H2O2 (1 mM) in artificial cerebrospinal fluid (aCSF) measured in the potential ranging from -1.2V to 0.8 V at a scan rate of 105 Vs-1. (b) Current-time response curve at rGO/Au2O3 electrode for successive injection of H2O2 (0.5 mM), AA (0.5 mM), DA (0.5 mM), glucose (1 mM), and UA (0.5 mM) in aCSF at -0.5 V.

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Figure 5. (a), (b) and (c) were the amperometric responses for rGO/Au2O3 electrode with deposition scan rates of 10, 50 and 250 mVs-1, respectively. (d), (e) and (f) were the corresponding calibration curves of the current versus H2O2 concentrations by amperometric measurement at applied potential of -0.50 V.

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Figure 6. (a) The amperometric response to H2O2 of rGO/Au2O3 electrode and non-coated electrode with detection limit at 1 µM and 100 µM at operation voltage of -0.5 V, respectively. (b) The calibration plot for rGO/Au2O3 electrode and non-coated electrode and their calibrations with the slope of 0.0038 nA‧µM-1 and 0.0025 nA‧µM-1 for rGO/Au2O3 electrode and non-coated electrode, respectively.

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Figure 7.

(a) The photomicrograph of a TTC- stained coronal section at 2 mm anterior to bregma.

The neural probe was implanted into the S1FL for real-time monitoring of the relative changes in electrochemical detection and real-time SSEPs recording (marked as black line). (b) Upper: The amperometric responses to H2O2 were recorded by non-coated gold electrode (Channel-2) and rGO/Au2O3 electrode (Channel-3) at -0.5V. At the time of 1500 sec, the S1FL brain was illuminated by laser light for 15-min to induce photothrombotic ischemic stroke (marked as square pale-yellow area). Lower: The first bare gold electrode (Channel-1) of the developed neural probe was used to record the changes in neuronal activities (SSEPs) before and after ischemic stroke.

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Acknowledgements This project was financially supported by the Ministry of Science and Technology the Republic of

China,

Taiwan

under

contract

numbers

of

MOST

102-2221-E-009 -024 -MY3.

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102-2221-E-010-011-MY3 and

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Supporting Information The Supporting Information is available free of charge on the ACS Publications website. SEM image and EDS analysis for rGO/Au2O3 of microelectrodes on neural probe, AFM image for rGO/Au2O3 of microelectrodes on neural probe. Raman spectral for rGO/Au2O3 of disposal gold electrode on plastic substrate (AUSPE, Zensor R&D, Taichung, Taiwan)

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Implantable Graphene-based Neural Electrode Interfaces for

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