In vitro and in vivo characterization of SU-8 flexible neuroprobe: From mechanical properties to electrophysiological recording

Sensors and Actuators A 216 (2014) 257–265

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Sensors and Actuators A: Physical journal homepage: www.elsevier.com/locate/sna

In vitro and in vivo characterization of SU-8 flexible neuroprobe: From mechanical properties to electrophysiological recording Shun-Ho Huang a , Shu-Ping Lin b,c , Jia-Jin Jason Chen a,d,e,∗ a

Department of Biomedical Engineering, National Cheng Kung University, Tainan, Taiwan Graduate Institute of Biomedical Engineering, National Chung Hsing University, Taichung, Taiwan Biomedical Engineering and System Bioinformatics Research Center, Kaohsiung Medical University, Taiwan d Medical Device Innovation Center (MDIC), National Cheng Kung University, Tainan, Taiwan e National Applied Research Laboratories, Taipei 106, Taiwan b c

a r t i c l e

i n f o

Article history: Received 3 March 2014 Received in revised form 12 May 2014 Accepted 6 June 2014 Available online 14 June 2014 Keywords: Neural implant Flexible neuroprobe SU-8 Biocompatibility Tissue reaction

a b s t r a c t Flexible neuroprobe with better tissue compliance exhibits unique mechanical characteristics in maintaining stability of neural implant in vivo. In this study, a flexible neuroprobe using SU-8 was designed and fabricated for in vitro and in vivo electrical sensing to show the improved tissue compatibility compared to that of the traditional rigid neuroprobe. The validation of neuroprobe was achieved by in vitro mechanical and cytotoxicity tests as well as in vivo neural recording and immunohistological staining. The fabrication process consisted of the creation of a backbone structure using photolithography, photopatterning of evaporated metal, and insulating of the electrode trace. The results of mechanical test of our fabricated SU-8 neuroprobe showed four times of physical stress (18.77 mN) than the insertion force (4.69 mN) to sustain resistance from brain tissue during implantation. The in vitro cytotoxicity assay showed well neuronal survival and proved the sufficient surface biocompatibility of the SU-8 neuroprobe. Further in vivo immunohistological staining showed no obvious glia aggregation around the implantation site indicating suitable biocompatibility compared with that of a rigid neuroprobe. Our in vitro and in vivo studies showed SU-8 neuroprobe possessed enough stress to complete the implantation in brain tissue and remained flexibility to comply micromovement of soft tissue with minor immune responses to achieve in vivo electrophysiological recordings at a signal-noise-ratio of greater than 7. © 2014 Elsevier B.V. All rights reserved.

1. Introduction The tissue compliance of an implanted neuroprobe is a critical issue for the short-, long-term stability and functionality of electrophysiological recording [1,2]. For probe strength or capacity of standard silicon fabrication process, implantable neuroprobe, such as the Utah microelectrode array and the Michigan probe, are made of rigid silicon-based materials, and could thus cause severe damage during inserting into the brain tissue [3]. Research has shown that the mechanical mismatch between a rigid silicon probe (whose Young’s modulus is about 170 GPa) and soft brain tissue (whose Young’s modulus is about 10 kPa) can induce chronic

∗ Corresponding author at: Department of Biomedical Engineering, National Cheng Kung University, No. 1 University Road, Tainan, Taiwan. Tel.: +886 6 2757575x63423; fax: +886 6 2343270. E-mail address: [email protected] (J.-J.J. Chen). http://dx.doi.org/10.1016/j.sna.2014.06.005 0924-4247/© 2014 Elsevier B.V. All rights reserved.

inflammation around the implantation site [4,5]. Furthermore, the changes in the mechanical properties of the surrounding environment cause specific tissue responses [6–8]. A glial sheath proliferated from glial cells such as astrocytes and microglia could encapsulate the probes and eventually isolate the electrodes from surrounding neurons after implantation [9,10]. Polymer materials with two ranges of mechanical properties have been used for fabricating flexible neuroprobe which include polydimethylsiloxane (PDMS) [11] with Young’s modulus of 1–10 MPa and polyimide (PI) [12], parylene [13] and SU-8 [38] with the range of 1–10 GPa. These polymer materials have better tissue compliance compared to the rigid material like silicon or glass with Young’s modulus ≥10 GPa [4,5]. In contrast with a silicon-based neuroprobe, a flexible PI probe can minimize the tissue response due to its lower strain force at the probe tip, as observed in finite element modeling [14] and in chronic tissue inflammation assay of animal study [15]. Flexible materials such as PI [16,17], parylene [18], and SU8 [19–22] have been adopted for the fabrication of neuroprobe

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Fig. 1. Schematic representation of the fabrication processes for SU-8 neuroprobe in oblique and cross section views. (a) E-beam evaporation of Al sacrificial layer on a glass substrate, (b) spin-coating and definition of SU-8 neuroprobe pattern, (c) E-beam evaporation of Ti/Au electrode layer, (d) spin-coating and definition of electrode trace and pad by photoresist (PR), (e) etching of unwanted metal using wet etching solution, (f) removal of photoresist via acetone wash, (g) spin-coating and definition of opening via SU-8 patterning, (h) release of complete SU-8 neuroprobe via etching of Al sacrificial layer, and (i) layout of SU-8 neuroprobe with two groups of microelectrode for two different depths of brain regions, i.e. cortex and striatum.

to improve biocompatibility after implantation. PI-based flexible neuroprobe was relatively easy to manipulate and fabricate in the conventional Micro-Electro-Mechanical-Systems (MEMS) industry because of some beneficial properties of PI, such as excellent insulation, dielectric strength, and mechanical flexibility [16,17]. In addition, surface modified PI substrate with biomolecule improves its biocompatibility [23]. However, PI neuroprobe complicate the surgical procedure due to insufficient mechanical strength [18,24]. Some mixtures of materials have been utilized to strengthen the mechanical property of relatively soft implantable neuroprobe. The stiffness of a flexible parlyne neuroprobe with fluid channels was enhanced using PEG coating for easy insertion into the cortex [18]. A layer of silicon or metal [25,26] has also been added to increase the stiffness of a PI neuroprobe; however, the combination of several processes might complicate the fabrication procedure. Another approach is to coat the probe with ultrafast resorbing or degradable polymers, which provide enough stiffness for implantation but are absorbed after immersion in the physiological environment [27,28]. Although PI and parylene based flexible neuroprobes showed their good tissue compatibility and electrophysiological recording functionality, the insufficient mechanical strength make the surgical operation difficult. On the other hand, SU-8 neuroprobe provide adequate stiffness and eliminate unnecessary modification for surgical implantation [20–22,29]. With the advantage of fabrication capacity of thick-film around hundred ␮m, the SU-8 neuroprobe represented suitable mechanical properties as backbone supportive material for flexible neural implant. The aims of this study are to fabricate a flexible SU-8-based neuroprobe for alleviating the surgical complication and reducing the immune response after implantation. We suggest that the SU-8 neuroprobe can not only sustain resistance from brain tissue during surgical implantation but also remain flexible to prevent chronic tissue damage of the surrounding area. The biocompatibility of the SU-8 neuroprobe can be assessed by in vitro cytotoxicity assay and immunohistological staining of the neurons and astrocytes around the implanted site. The proposed neuroprobe can be used to record neural activity without causing severe tissue damage.

2. Materials and methods 2.1. Design and fabrication process of SU-8 neuroprobe The flexible neuroprobe comprised of a multilayer of SU-8 using conventional MEMS technique with a releasing method of sacrificial layer. The fabrication process of the SU-8 neuroprobe began with a thin layer of aluminum as a sacrificial layer on a glass substrate, as depicted in Fig. 1(a). The first layer of SU-8 (80 ␮m) was spun and soft-baked at 65 ◦ C and then at 75 ◦ C. Photolithography was then applied to define the backbone structure of the SU-8 neuroprobe, followed by post-baking, as depicted in Fig. 1(b). Once this layer was developed, the thermal evaporation of titanium (20 nm) and gold (200 nm) was performed using an E-beam (Fig. 1(c)). In order to define the electrode for electrophysiological recording, a positive photoresist (S1818) was photolithographically patterned as the cover layer to avoid unwanted etching (Fig. 1(d)). After the process of wet etching, Fig. 1(e) showed the remaining probe substance of electrodes, pads and circuits. Acetone washing was then conducted to remove the S1818 (Fig. 1(f)). Another 6␮m SU-8 film was then spun and patterned as the insulating layer of the metal traces (Fig. 1(g)). Finally, the complete structure of the SU-8 neuroprobe was released by the removal of the sacrificial layer, as shown in Fig. 1(h). Fig. 1(i) showed the layout of the SU8 neuroprobe. The appearance of our proposed SU-8 neuroprobe was further examined by scanning electron microscopy (SEM). The impedance properties of microelectrodes of SU-8 neuroprobe were then characterized using a precision LCR meter (Agilent 4284A, Agilent Technologies) under a constant 10-mVrms AC voltage at 1 kHz. The acquired data were displayed in real-time in the LabVIEW graphical user interface (GUI) and saved for off-line analysis. 2.2. Evaluation of mechanical properties of SU-8 neuroprobe A previously reported analytical model of bending and buckling force for evaluating the mechanical strength needed to penetrate brain tissue was adopted [30]. Two main forces, the bending force and the buckling force, were measured during implantation. The

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bending force resulted from an out-of-plane loading that caused displacement parallel to the tissue plane. The buckling force was from uniaxial compression and represented the counteracting force exerted normally on the brain tissue by the probe and the shear force preventing the tip of the probe from slipping on the surface of the brain tissue. A tensile tester was utilized for measuring the maximum force of bending and buckling until the probe broke. A simple brain phantom with a reproducible elastic modulus of 10 kPa equivalent to the brain was composed of a concentration of 0.6% agar gel in deionized water and adopted for the insertion test [31,32]. A custom-made tensile tester was used for measuring the force changes during insertion into and retraction from the phantom. A load cell force sensor (2 N, with a resolution of 0.2 mN) combined with a motorized motion controller at a resolution of ∼10 ␮m was employed. The probe was carefully aligned in the vertical position to ensure the reproducibility of experimental results. In addition to the agar phantom, biological rat brain tissue was held in an ice-cold glass dish for measuring the force response during the insertion process. The force–time curves during the insertion and retraction of the probe were calculated for determining the mechanical strength of the SU-8 neuroprobe. 2.3. In vitro viability test and extracellular recording using neuroprobe Primary cortical neurons were prepared as described previously [33] and were used to culture on the SU-8 neuroprobe for the in vitro biocompatibility tests. After 7 and 14 days in vitro (DIVs), the Live/Dead Viability/Cytotoxicity Kit (L3224, Molecular Probes, Eugene, OR) was utilized for quantifying cell viability. Fluorescence images were acquired using a charge-coupled device (CCD, CoolSNAP, Media Cybernetics, Silver Spring, MD) attached to an inverted fluorescence microscope (IX71, Olympus, Tokyo, Japan). The numbers of live and dead cells were calculated using image processing software (ImageJ, National Institutes of Health). The percentage of viable cells was calculated by dividing the number of live cells by the number of total cells. The extracellular recording of in vitro neuronal networks was performed after 14 DIVs and carried out using the gold microelectrode of the SU-8-based neuroprobe. A platinum wire was immersed in the same medium as a reference electrode. The signals from the electrodes were amplified (1000×) and band-pass filtered (300–3000 Hz) with a pre-amplifier (MultiChannel Systems, MPA8I) and a programmable amplifier (MultiChannel Systems, PGA32). Signals of all channels were simultaneously sampled at 20 kHz per channel by a data acquisition device (National Instruments, Austin, TX, PCI-6052E) which was controlled by GUI programmed in LabVIEW software.

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area (above the ventral striatum, 1.5 mm anterior and 1.8 mm lateral to the bregma) [34]. The neuroprobe was penetrated through the pia into the brain until it was 5.0 mm below the brain surface. The flexible neuroprobe with a PCB was attached to the skull and encased to seal the craniotomy and the connector assembly using dental acrylic. A stainless steel ground wire attached to the preamplifier was connected to a bone screw to provide a ground point. The electrophysiological recording started 30 min after SU-8 neuroprobe implantation. The animal was anesthetized during the entire recording period. The neural signals were sensed from the gold microelectrode of the SU-8 neuroprobe with a stainless steel wire as the reference electrode. The signal acquisition hardware was the same as that used for in vitro extracellular recording. Data were analyzed offline using MATLAB (The MathWorks Inc., USA) based on a typical spike sorting algorithm to estimate the signal-to-noise ratio (SNR) [35]. Sensing noise, instrumentation interference, and small amplitude signals from neighboring neurons were considered as background noise. An amplitude threshold based on three times and the standard deviation was used for distinguishing spike signals from background noise. For investigating the acute and chronic immune response of tissue reaction to the implant, in vivo biocompatible testing was performed via the immunohistological technique for staining neurons and glia. Three animals were sacrificed for immunohistological staining. After implantation, the brain was removed at either 30 min or 2 weeks after implantation. The animals were anesthetized with an over-dose of chloride hydrate of IP injection and transcardially perfused with 150–200 ml of ice-cold 0.1 M phosphate-buffered saline (PBS), followed by the fixation of 4% paraformaldehyde in 0.1 M PBS at 4 ◦ C for 1 day. Samples were soaked for 2 days in 0.1 M PBS containing 30% sucrose for preservation. A cryosection was subsequently performed in the horizontal plane at 40-␮m resolution using a sliding knife freezing microtome. Perpendicular alignment of the brain could help to define the depth of the probe and the corresponding section area. The primary antibodies were rabbit polyclonal antibodies recognizing glia fibrillary acidic protein (GFAP, an astrocytic cytoskeleton protein, 1:5000, Chemicon, USA) and mouse monoclonal antibodies recognizing MAP II (neuronal marker, 1:100, Chemicon, USA). The secondary antibodies were Alexa488-conjugated antibodies for mouse IgG and Alexa594-conjugated antibodies for rabbit IgG (1:100, Invitrogen, USA). Following rinses of the section in PBS and pre-incubation in a mixture of 5% normal goat serum and 0.25% Triton X-100 (Sigma, Germany) in 0.02 M PBS, samples were reacted with the primary antibodies over-night at room temperature. After several rinses in PBS, samples were then incubated with secondary antibodies in a dark environment for 2 h at room temperature and then rinsed in PBS. Finally, samples were preserved in a microscope slide for imaging.

2.4. In vivo electrophysiological recording and immunobiological assays 3. Results All animals were obtained from the Animal Center of Medical College, National Cheng Kung University, Taiwan, and all surgical procedures were performed according to the Guide for the Care and Use of Laboratory Animals. SU-8 neuroprobe was surgically implanted in male Sprague Dawley rats (300–400 g). The animals were administered general anesthesia reagents, including a mixture of 50 mg/ml ketamine, 5 mg/ml xylazine, and 1 mg/ml acepromazine. The anesthesia reagents were intraperitoneally injected at an initial dosage of 0.1 ml/100 g body weight and maintained in an areflexive state throughout the surgical procedure using regular anesthesia supplements with 20% initial dose. The animal was fixed in a standard stereotaxic frame, and three stainless steel bone screws were then inserted into its skull. A craniotomy (about 1 mm × 1 mm) was made over the target brain

3.1. Fabrication and electrical characterization of SU-8 neuroprobe The total length of the SU-8 neuroprobe was 7 mm. Four recording electrodes were divided into two groups, and then each group was used for simultaneously monitoring different depths of brain regions, i.e. cortex and striatum. Fig. 2 shows the SEM images of four electrodes (diameter = 50 ␮m, circular shape) for electrophysiological signal recording and one electrode reserved for integrating electrochemical sensing in the future (dimensions = 30 ␮m × 100 ␮m, rectangular shape). The electrodes were immersed in PBS and characterized by an impedance analyzer. The measured microelectrode impedance was 150 ± 30 k at 1 kHz,

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Fig. 2. SEM images of SU-8 neuroprobe of (a) whole probe structure and (b) one group of microelectrodes with four microelectrodes for electrophysiological recording (circular shape) and one reserved for future integration of electrochemical sensing (rectangular shape). Higher magnification of (c) single recording microelectrode and (d) wire bonding to external PCB for recordings.

which is suitable for neuronal recording. The low impedance value of the microelectrode is due to its large surface area compared to that of Michigan probe [35]. The long-term stability of the SU-8 neuroprobe was confirmed by impedance measurements in PBS over 14 consecutive days. The impedance values of the microelectrodes were 245 ± 17 k before soaking and 251 ± 23 k after 14 days soaking showing no significant difference (N = 40) under Student’s t-test. 3.2. Mechanical characterization of SU-8 neuroprobe Fig. 3(a) illustrates the deflection of the neuroprobe during the insertion phase. As shown in Fig. 3(b), the critical buckling force is an essential factor for the assessment of probe failure. Here, we utilized a tensile tester to measure the maximum value of directional buckling and bending force. For the SU-8 neuroprobe, the maximum buckling force was 18.77 mN, which indicates sufficient mechanical strength for surgical implantation. Bending test showed that the SU-8 neuroprobe possesses substantial flexibility and can be bent to more than 60◦ without probe failure (Fig. 3(c)). Fig. 4 shows the real-time force measurement during the insertion into and retraction during three samples: a phantom (0.6% agar gel) and a rat’s brain with and without dura. All experiments were performed at fixed low speed (100 ␮m/s) for insertion and retraction which mimic the speed of the surgical implantation in vivo. The SU-8 neuroprobe penetrated the phantom and the rat’s brain without any failure for all experiments (N = 2 for each sample). The SU-8 neuroprobe was inserted into the brain phantom smoothly, with a plateau resistance force of about 2.05 mN. The rat’s brain samples

containing pia/dura or pia showed a larger insertion force in comparison with agar phantom. With the sample of brain with dura which has higher elastic modulus of dura matter compared to agar phantom, the probe structure bent with a measured peak force of 4.69 mN, the force then rapidly dropped to around 3 mN after the dura had been penetrated. The measured force from brain without dura showed an increased trend with fluctuated pattern and reached around 3 mN after insertion. The results show that our proposed neuroprobe are strong enough to withstand the resistance produced by the phantom and brain tissue. 3.3. In vitro cytotoxicity testing assay and extracellular physiological recording A cytotoxicity test was performed using the primary cortical neuronal culture. Fig. 5 shows the cell viability and fluorescence images of cortical neurons on culture dish and SU-8 neuroprobe for 7 and 14 DIVs. The formation of a neuronal network consisting of healthy neurons and neurite extensions was observed. From the cell viability assay, the survival rate of neurons on the SU-8 probe was comparable to that on culture dish at 7 DIVs (culture dish: 74.3 ± 6.2%, SU-8 neuroprobe: 84.3 ± 3.5%, N = 10 for each group). Obvious neurite outgrowth on the surface of SU-8 neuroprobe was observed. For a longer culture of 14 DIVs, the cell viability (culture dish: 76.8 ± 4.9%, SU-8 neuroprobe: 79.9 ± 4.8%) showed sufficient cell survival rate on culture dish and SU-8 neuroprobe. Although the cell viability seems be lower on culture dish, no statistically significant difference was observed. Slight mean cell viability difference might due to the difference in surface properties between petri-dish and SU-8 [36,37].

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Fig. 3. Mechanical evaluation of SU-8 neuroprobe. (a) Mechanical analytical model of SU-8 neuroprobe during insertion process and images acquired during, (b) buckling test and (c) bending test.

Extracellular physiological recording was performed on 17 DIVs of neuronal culture while the neuronal network was well developed on SU-8 neuroprobe. Four experiments were successfully performed with spontaneous activities pattern. Fig. 6 shows that the spontaneous action potential of cortical neurons was clearly recorded by the microelectrodes of SU-8 neuroprobe. In addition, several concurrently spontaneous activities were observed in 30 s. During the recording from SU-8 neuroprobe (N = 4), 20 of the 32 electrodes were able to record clear spontaneous activity with a 62.5% success rate. 3.4. In vivo immune response and electrophysiological recording After the implantation of SU-8 neuroprobe, the immune response was thus tested using anti-MAPII (neurons) and anti-GFAP (astrocytes) for visualization of neuron-astrocyte interaction. Two animals were sacrificed for staining. As shown in Fig. 7, no apparent

astrocyte aggregation was observed in the cortex region (lesion width: 400 ␮m) and striatum (lesion width: 200 ␮m) after 2 weeks of implantation. Fig. 7(a) shows a thin layer of astrocyte aggregation around the surrounding of lesion site, but without the formation of the glia sheath. Similar result was observed in the lesion area of striatum (Fig. 7(b)). The evident visualization of neurons located adjacent to the electrodes revealed the long-term stability of the SU-8 neuroprobe. For acute recording, five experiments of SU-8 neuroprobe implantation were performed. Fig. 8 shows that neural signals were recorded 30 min after the SU-8 neuroprobe was implanted into the rat’s brain under anesthesia. The maximum amplitudes are around 628 ␮V (peak-to-peak value), with the background noise typically below 40 ␮Vrms after the removal of the power line interference. From two examples of lowest and highest motor unit activities, the SNR values were 7.9 and 23.9, respectively, as shown in Fig. 8(b) and (c).

Fig. 4. Insertion and retraction behavior of flexible SU-8 neuroprobe during three different samples. The insertion resistance force exhibited different patterns for 0.6% agar brain phantom (solid line), brain with dura and pia (dotted line), and brain with pia only (dashed line). The insertion of neuroprobe was operated at fixed speed of 100 ␮m/s to around 6 mm in depth. Before retraction at 80 s, the resistance force dropped to about zero after insertion.

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Fig. 5. Cytoxicity test of SU-8 neuroprobe. Cortical neurons cultured in culture dish for (a) 7 DIVs and (c) 14 DIVs. Cortical neurons cultured on SU-8 neuroprobe for (b) 7 DIVs and (d) 14 DIVs. (e) Quantification of cell viability for culture and neuroprobe (N = 10 for each group).

4. Discussion SU-8 has been commonly used as a structuring material in MEMS fabrication due to its easy process capability [38,39] which also provides adequate stiffness of Young’s modulus about 4 GPa for

flexible neuroprobe design [38]. This study demonstrated the fabrication of a flexible SU-8 neuroprobe with good mechanical and biocompatible properties. Experiments were conducted to investigate the mechanical properties of our flexible SU-8 neuroprobe. Two important mechanical properties regarding the critical force

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Fig. 6. Extracellular spontaneous action potential of cortical neurons cultured on SU-8 neuroprobe was recorded from 17-DIV cortical neurons, (a). Magnified view of signals recorded from SU-8 neuroprobe can be distinguished from background noise, (b).

for probe failure and the on-line force measurement during surgical implantation were respectively characterized. Flexible materials with tunable softness (Young’s modulus = 1–10 MPa for elastomers, 1–10 GPa for plastics) have been used for fabricating varied neural implants with better tissue compliance in contrast with a rigid material like silicon or glass (Young’s modulus  10 GPa) [4,5]. However, relatively soft materials cause difficulty in handling neural implants during implantation. Several approaches have been used to tackle this problem, such as a partly attached silicon backbone [40,41] which served as supporting material for overcome the surgical complication or additional coating/filling materials [18,42] to enforce the stiffness of a flexible probe. Some mechanical adaptive materials represented excellent properties which have mechanical sufficient stiffness during surgical implantation then soften in vivo could be potential candidates as flexible neuroprobe

backbone material [43,44]. Our SU-8 neuroprobe provides suitable stiffness during surgical insertion without any redundant preparation. From the mechanical evaluation of the SU-8 neuroprobe, the maximum buckling force (18.77 mN) is four-time larger than the peak insertion force measured in the brain sample with dura and pia (4.69 mN) (Fig. 4). Although the maximum buckling force value is lower than that of a rigid material [45], the sufficient buckling force resistance of our SU-8 neuroprobe makes the implantation procedure easier. These results demonstrated that the SU-8 neuroprobe can withstand the force during penetration into brain tissue. Biocompatibility and biological response to the implanted neuroprobe are the major challenges to maintain stable neural implantation [1,2]. The mechanical trauma during insertion, foreign body reaction, implantation method, and physical properties of the electrodes and backbone material further affects the

Fig. 7. Immunohistological staining of tissue reaction around implantation site. Anti-MAPII/GFAP staining around implanted neuroprobe in (a) cortex region and (b) striatum 14 days after implantation (anti-MAPII stained red for neuron marker; anti-GFAP stained green for astrocyte marker).

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Fig. 8. (a) Electrophysiological recording of neural signals after SU-8 neuroprobe implantation. Two action potential signals extracted from those marked with arrows in (a) with, (b) minimum and (c) maximum SNR.

biocompatibility and biological response [2]. Factors such as surface biocompatibility, structure biocompatibility, and biostability should be considered when developing a neuroprobe [5]. This study showed that the surface biocompatibility of SU-8 is comparable to that of a conventional Petri dish according to the relatively high viability of cortical neurons in Fig. 5. The cell viability on SU-8 neuroprobe has a comparable survival rate at 7 and 14 DIVs and showed well formation of in vitro neuronal network. Our results of cell viability also indicated the non-toxicity of SU-8 [37]. The tissue response to the implanted SU-8 neuroprobe was verified by anti-MAPII (neurons) and anti-GFAP (astrocytes), which have been commonly used for the visualization of neuron–astrocyte interaction with implanted devices [46,47]. The immunohistological staining of MAPII/GFAP for neuron distribution and reactive gliosis was clearly visualized around SU-8 neuroprobe and showed less surgical damage and chronic inflammation in comparison with rigid silicon probes in other studies [9,10,46,47]. Fig. 7(a) and (b) shows that no clear glia sheath formed after SU-8 neuroprobe implantation for 14 days, either in the cortex (larger lesion area) or striatum (smaller lesion area) region. In addition, the lesion area in the striatum region had less astrocyte aggregation and more neurons in contact with the implantation site compared with those in the cortex region. The main drawback of rigid silicon probes is the mismatch between their mechanical properties and brain tissue and eventually leads to cellular reactions and encapsulation of the implant [48]. The SU-8 neuroprobe has a thin-film-like shape, which prevents surgical damage during implantation [9], and high flexibility, which reduces tissue reaction to the implanted device [19]. Moreover, SU-8 has been previously shown good

biocompatibility in vitro and in vivo [36,37] and suitable mechanical properties as a flexible neural implant [5]. In addition, our results of immunohistological staining showed biocompatible SU-8 neuroprobe can reduce immune response after implantation. The electrophysiological recordings were performed in an in vitro cortical neuronal network and in vivo implantation, respectively. The spontaneous action potentials of in vitro cultured neurons on SU-8 neuroprobe provided useful information of cytotoxicity and functionality [49,50]. The spontaneous activities of cortical neurons recorded at 17 DIVs demonstrated the feasibility of the developed SU-8 neuroprobe for recording neuronal activities. After implantation, the results of electrophysiological recording (Fig. 8) show several spikes in 45 s with a SNR ranging from 7.9 to 23.9. In a 45-s section of neuroprobe recordings, multiple action potential of neuronal activities were recorded simultaneously in contrast to the action potential recorded in vitro (Fig. 6). The SNR values are comparable to those for a silicon probe [35] or a wire electrode [51]. The flexibility of our SU-8 neuroprobe provides substantial biocompatibility and tissue compliance to prevent tissue damage caused by micro-motion. In addition, in vivo electrophysiological recording demonstrated the possibility of utilizing SU-8 neuroprobe to practically record neural activities in living animals. Previous works based on development of SU-8 neuroprobe have demonstrated the electrophysiological recordings from peripheral nerve and hippocampus [21,22,29] which showed the SU-8 a potential candidate as backbone material of flexible neuroprobe. However, there are still lack of validation processes of SU-8 based on mechanical properties and biocompatibility. In this study, we performed mechanical evaluation of SU-8 neuroprobe and showed

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sufficient buckling force resistance for surgical implantation. In addition, our results of in vitro cytotoxicity and immunohistological staining of neurons and astrocytes further showed the suitable biocompatibility of SU-8. It is known that the biocompatibility of implanted device could be affected by multiple factors such as mechanical trauma during insertion, foreign body reactions of inflammation and chronic neurodegeneration, as well as mechanical mismatch between implanted neuroprobe and micro-motion ([1,2,4,5]. Among these factors, a flexible neuroprobe can mainly reduce the mechanical mismatch which mimics mechanical properties to the brain tissue compared to the rigid neuroprobe. Our results provide mechanical and biocompatibility evidences for fabricating flexible neuroprobe based on SU-8. 5. Conclusion In this study, SU-8 flexible neuroprobe was fabricated for reliable electrophysiological recording. Mechanical characterization of the SU-8 neuroprobe revealed that it has good mechanical strength and flexibility. Qualitative results from in vitro cytotoxicity tests and in vivo immunohistological staining show that SU-8 neuroprobe owns suitable properties for neural implant applications. Our on-going project is to integrate the SU-8 neuroprobe with various recording or stimulation strategies, such as electrochemical sensing [52], optical stimulation [53], or wireless recording [54], to create a neural implant with multiple modalities. Acknowledgements This work was supported by the National Science Council of Taiwan under grant NSC 101-2221-E-006-006-MY3 and National Health Research Institutes of Taiwan under grant NHRI-EX10210139EI. Instrumental support for fabricating the SU-8 neuroprobe was given by the Micro-Nano Technology Research Center and the Southern Region Micro-Electro-Mechanical Systems Research Center in Taiwan.

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[29] S.H. Cho, H.M. Lu, et al., Sens. J. IEEE 8 (2008) 1830–1836. [30] M. Hajj-Hassan, V. Chodavarapu, et al., J. Micro/Nanolithogr. 8 (3) (2009) 033011. [31] T.J. Hall, M. Bilgen, et al., IEEE Trans. Ultrason. Ferroelectr. Freq. Contr. 44 (6) (1997) 1355–1365. [32] R. Hoffmann, T. Stieglitz, et al., IEEE Annu. Conf. Eng. Med. Biol. Soc. (2007) 4711–4714. [33] S.H. Huang, S.P. Lin, et al., Biomed. Microdev. 15 (2012) 135–143. [34] G. Paxinos, C. Watson, The Rat Brain in Stereotaxic Coordinates, Academic Press, Oxford, UK, 2007. [35] R.J. Vetter, J.C. Williams, et al., IEEE Trans. Biomed. Eng. 51 (6) (2004) 896–904. [36] G. Voskerician, M.S. Shive, et al., Biomaterials 24 (11) (2003) 1959–1967. [37] V.N. Vernekar, D.K. Cullen, et al., J. Biomed. Mater. Res. A 89 (1) (2009) 138–151. [38] H. Lorenz, M. Despont, et al., J. Micromech. Microeng. 7 (3) (1997) 121. [39] P. Abgrall, V. Conedera, et al., Electrophoresis 28 (24) (2007) 4539–4551. [40] K. Lee, A. Singh, et al., Sens. Actuators B: Chem. 102 (1) (2004) 67–72. [41] E.G.R. Kim, J.K. John, et al., Sens. Actuators B: Chem. 195 (2014) 416–422. [42] L.W. Tien, F. Wu, et al., Adv. Funct. Mater. 23 (2013) 3185–3193. [43] J.P. Harris, A.E. Hess, et al., J. Neural Eng. 8 (2011) 046010. [44] T. Ware, D.D.E. Simon, et al., Adv. Funct. Mater. 22 (2012) 3470–3479. [45] K. Najafi, J. Hetke, IEEE Trans. Biomed. Eng. 37 (5) (1990) 474–481. [46] J. Turner, W. Shain, et al., Exp. Neurol. 156 (1) (1999) 33–49. [47] V.S. Polikov, P.A. Tresco, et al., J. Neurosci. Methods 148 (1) (2005) 1–18. [48] D.F. Williams, Biomaterials 29 (20) (2008) 2941–2953. [49] S. Martinoia, L. Bonzano, et al., Biosens. Bioelectron. 20 (10) (2005) 2071–2078. [50] G. Xiang, L. Pan, et al., Biosens. Bioelectron. 22 (11) (2007) 2478–2484. [51] J.C. Williams, R.L. Rennaker, et al., Brain Res. Protoc. 4 (3) (1999) 303–313. [52] M.D. Johnson, R.K. Franklin, et al., J. Neurosci. Methods 174 (1) (2008) 62. [53] A.V. Kravitz, A.C. Kreitzer, Curr. Opin. Neurobiol. 21 (3) (2011) 433–439. [54] I.H. Hsieh, C.C.-H. Yang, et al., J. Med. Biol. Eng. 33 (2013) 79–86.

Biographies

Shun-Ho Huang received the B.S. degree from National Kaohsiung University of Applied Sciences, Kaohsiung, Taiwan, in 2006, and the M.S. and Ph.D. degrees from National Cheng Kung University, Tainan, Taiwan, in 2008 and 2014, respectively, all in biomedical engineering. His present research activities involve microelectrode arraysbased electrophysiological characterization for in vitro/in vivo 3D neuronal network via impedimetric sensing and extracellular and pathological environmental simulation for in vitro ischemic stroke model and in vitro Alzheimer’s disease model.

References [1] [2] [3] [4] [5] [6] [7] [8] [9] [10] [11] [12] [13] [14] [15] [16] [17] [18] [19] [20] [21] [22] [23] [24] [25] [26] [27] [28]

J.B. Leach, A.K.H. Achyuta, et al., Front. Neuroeng. 2 (2009). C. Marin, E. Fernandez, Front. Neuroeng. 3 (2010). M.P. Ward, P. Rajdev, et al., Brain Res. 1282 (2009) 183–200. S.P. Lacour, S. Benmerah, et al., Med. Biol. Eng. Comput. 48 (10) (2010) 945–954. C. Hassler, T. Boretius, et al., J. Polym. Sci. B: Polym. Phys. 49 (1) (2011) 18–33. D.E. Discher, P. Janmey, et al., Science 310 (5751) (2005) 1139. F. Rehfeldt, A.J. Engler, et al., Adv. Drug Deliv. Rev. 59 (13) (2007) 1329–1339. F.C. Su, C.C. Wu, et al., J. Med. Biol. Eng. 31 (2011) 233–244. D. Szarowski, M. Andersen, et al., Brain Res. 983 (1) (2003) 23–35. R. Biran, D.C. Martin, et al., Exp. Neurol. 195 (1) (2005) 115–126. J.C. Lotters, W. Olthuis, et al., J. Micromech. Microeng. 7 (1997) 145. R. Djakaria, B.I. Chandran, et al., 47th Proceedings of Electronic Components and Technology Conference, 1997, pp. 745–749. H.S. Noh, K.S. Moon, et al., J. Micromech. Microeng. 14 (2004) 625. J. Subbaroyan, D.C. Martin, et al., J. Neural Eng. 2 (2005) 103. K. Cheung, Y. Zhong, et al., Eur. Cells Mater. 10 (2005) 5. K.C. Cheung, P. Renaud, et al., Biosens. Bioelectron. 22 (8) (2007) 1783–1790. A. Mercanzini, K. Cheung, et al., Sens. Actuators A: Phys. 143 (1) (2008) 90–96. S. Takeuchi, D. Ziegler, et al., Lab Chip 5 (5) (2005) 519–523. M. Tijero, G. Gabriel, et al., Biosens. Bioelectron. 24 (8) (2009) 2410–2416. A. Altuna, G. Gabriel, et al., J. Micromech. Microeng. 20 (6) (2010) 064014. A. Altuna, L.M. de la Prida, et al., Biosens. Bioelectron. 37 (1) (2012) 1–5. A. Altuna, E. Bellistri, et al., Lab Chip 13 (2013) 1422–1430. J.M. Goddard, J. Hotchkiss, Prog. Polym. Sci. 32 (7) (2007) 698–725. T.D.Y. Kozai, D.R. Kipke, J. Neurosci. Methods 184 (2) (2009) 199–205. K.K. Lee, J. He, et al., J. Micromech. Microeng. 14 (1) (2003) 32. Y.Y. Chen, H.Y. Lai, et al., J. Neurosci. Methods 182 (1) (2009) 6–16. D. Lewitus, K.L. Smith, et al., Acta Biomater. 7 (6) (2011) 2483–2491. D.Y. Lewitus, K.L. Smith, et al., Biomaterials 32 (24) (2011) 5543–5550.

Shu-Ping Lin obtained her Ph.D. in Biomedical Engineering from National Cheng Kung University in 2008. She was a visiting research associate in Dr. Themis Kyriakides’s laboratory at Yale University from 2006 to 2007. She worked as a postdoctoral researcher at the Institute of Atomic and Molecular Sciences, Academia Sinica (2008-2009), and independent researcher at Industrial Technology Research Institute (2009-2010). In 2010, she joined Graduate Institute of Biomedical Engineering, National Chung Hsing University, as an assistant professor. She has published papers in reputed journals, such as NanoToday, Biomaterials, Biomacromolecules, Biomedical Microdevices, and Sensors, etc. Her research interests are mainly directed toward designing and developing biocompatible nano/micro biosensors, cell/tissue engineering, electrophysiological measurements, biocompatible and functionable surface modification, and characterization of biointerface. Jia-Jin Jason Chen received the B.S. degree from Chung Yuan Christian University, Chung-Li, Taiwan, in 1980, and the M.S. and Ph.D. degrees from Vanderbilt Uni- versity, Nashville, TN, in 1987 and 1990, respectively, all in biomedical engineering. Since 1997, he has been a Professor at Institute of Biomedical Engineering, National Cheng Kung University, Tainan, Taiwan. His present research activities involve biomedical signal processing, neural engineering, functional electrical stimulation, neural/neuronal interfaces, and implantable biomicrosystem.