Carbon Microfiber Electrodes

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Biofunctionalized conducting polymer / carbon microfiber electrodes for ultrasensitive neural recordings Hugo Vara, and Jorge Collazos-Castro ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.5b09594 • Publication Date (Web): 17 Nov 2015 Downloaded from http://pubs.acs.org on November 24, 2015

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ACS Applied Materials & Interfaces

Biofunctionalized conducting polymer / carbon microfiber electrodes for ultrasensitive neural recordings

Hugo Vara and Jorge E. Collazos-Castro*

Neural Repair and Biomaterials Laboratory, Hospital Nacional de Parapléjicos (SESCAM). Finca la Peraleda s/n, 45071 Toledo, Spain.

*

Corresponding author at:

Neural Repair and Biomaterials Laboratory, Hospital Nacional de Parapléjicos (SESCAM). Finca La Peraleda s/n, 45071 Toledo, Spain Tel.: +34 925247758 Fax: +34 925247745 E-mail: [email protected] ACS Paragon Plus Environment


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Abstract

Carbon microfibers (MFs) coated with conducting polymers may provide a solution for long-term recording of activity from individual or small groups of neurons. Attaching cell adhesion molecules to the electro-sensitive surface might further improve electrode-neuron contact, thus enhancing signal stability and fidelity. We fabricated biofunctionalized microelectrodes consisting of 7-µm diameter carbon MFs coated with poly(3,4ethylenedioxythiophene) doped with poly[(4-styrenesulfonic acid)-co-(maleic acid)] (PEDOT:PSS-co-MA), and linked N-Cadherin to the polymer surface. These electrodes were tested for recording artificially generated electric potentials, as well as multi-unit activity (MUA), sharp wave-ripple complexes (SWRs) and field excitatory postsynaptic potentials (fEPSPs) in rat hippocampal slices. The effects of electrode length and functionalization were compared. PEDOT:PSS-co-MA coating improved electric current detection and reduced the electrical noise but had no significant effect on the amplitude of recorded biopotentials. Surface biofunctionalization lowered the electric current flow, and further reduced the electrical noise. Additionally, it increased the amplitude of the recorded MUA, finally doubling the signal-to-noise ratio achieved with bare carbon MFs. Biofunctionalization benefits on MUA were apparent only for potentials from cells putatively adjacent to the microelectrode. Analysis of fEPSPs excluded adverse effects of functionalized electrodes in basal synaptic transmission. These results demonstrate the possibility of enhancing the amplitude and signal-to-noise ratio of neural recordings by coating the microelectrodes with conducting polymers modified with neural cell adhesion molecules, and support the use of biofunctionalized MFs in advanced neuroprosthetic devices.

Keywords: Microfiber, conducting polymer, PEDOT, N-Cadherin, biofunctionalization, neural recording, signal-to-noise ratio.

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Introduction Design of advanced recording neuroprostheses for chronic implanted use requires optimization of signal detection capabilities while improving electrode integration within the neural tissue. Electrode implantation inside the central nervous system evokes an inflammatory response1,2, accompanied by neural damage and followed by the formation of a fibrotic scar and/or tissue cavities around the electrode1,3,4. Fibrous encapsulation separates the electrode from electroactive cells and also increases the electrical impedance of the bio-electronic interface3. Furthermore, the chronic presence of an active electrode has been associated to neurodegeneration in its close proximity5. All these processes impair the electrode performance in the long-term. Electrode miniaturization is the first approach to overcome tissue damage concerns and simultaneously assure optimal recording resolution at the neuronal level. In this regard, the small size of carbon microfibers (CFs), together with their high electrical conductivity, tensile strength and flexibility, make them an excellent choice for electrophysiological applications6. CF microelectrodes perform as good as the best tungsten electrodes at recording extracellular activity7,8. In fact, carbon-fiber electrode arrays have been successfully used for long-term recording of small clusters of neurons in birds9. However, the uninsulated electrode tips were still relatively long (~ 90 µm), which accounted for a scarcity of isolated single-unit recordings within the multi-unit activity. Smaller CF tips could improve single-unit isolation, but the beneficial effects of reducing electrode size are counteracted by the increased noise level10. Hence, electrode miniaturization needs additional modifications to improve signal-recording resolution. For a constant electrode gross geometric area, a higher effective area (through surface micro/nanostructuring) favors signal-recording resolution by diminishing electric noise11. In this direction, carbon nanotubes have been used to confer better neural recording properties to tungsten electrodes12. Micro/nanostructuring of electrodes may also be achieved by coating them with conducting polymers (CP), which increases the effective electrode surface and reduces its electrical impedance by 2-3 orders of magnitude13,14. This approach improves in vivo neural recordings with ultrasmall electrodes. A relevant work in this area used poly(3,4ethylenedioxythiophene) doped with polystyrene sulfonate (PEDOT:PSS), electrodeposited at the very tip of 7-µm diameter CFs. Those microfiber-based electrodes were implanted into the rat cerebral cortex and allowed the detection of single-unit activity for up 1-2 ACS Paragon Plus Environment

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months, with attenuated tissue damage and scarring response15. More recently, CF/CP multi electrode arrays were fabricated16, paving the way for neuroprosthetic devices that have the potential to record from all active neurons in a given cortical layer. Despite these advances, electroconducting microfibers still need optimisation to provide single-unit recordings for years after implantation. This might be achieved by functionalizing the CP with biomolecules that promote a direct contact between electrode and neural cells, thus avoiding the attenuation of the signal along the extracellular fluid17,18. Studies from our laboratory19 showed that cultured cerebral cortex neurons (CNs) could grow for several weeks on PEDOT:PSS films functionalized with polycations (polylysine and spermine), linked by electrostatic interactions. We also developed a more stable functionalization by electrosynthesizing poly[(4-styrenesulfonic acid)-co-(maleic acid)]doped PEDOT (PEDOT:PSS-co-MA) on metallic electrodes or carbon microfibers, and then covalently bonding polylysine or antibodies that linked recombinant neural cell adhesion molecules (L1 and/or N-Cadherin)20. Such functionalization allowed extensive dendrite and axonal growth of cultured CNs. The present work tests the hypothesis that functionalization with N-Cadherin may improve the quality of the electric signals recorded with CP-coated carbon microfibers. We describe the methodology necessary to obtain fully biofunctionalized electrodes, and characterize their electrical features in the different steps of fabrication. Since no systematic investigations have been performed to date with such CP/CF electrodes, several electrode lengths are evaluated in artificial electrolytes and also with neural tissue, in order to inform size optimization. The general electrode properties are studied by cyclic voltammetry and chronoamperometry, whereas the electro-sensing performance is investigated in response to externally applied voltage pulses. Then, the actual applicability of the electrodes for neural activity recording is assessed in the acute in vitro rat hippocampal slice. This includes the analysis of electric potentials generated by neurons close or in contact to the electrode surface, and likewise potentials arising from the activity of neuronal networks distributed within the tissue. The results provide evidence that N-Cadherin biofunctionalized CP/CF electrodes display advantageous properties for the selective detection of the activity of neighboring neurons. Furthermore, measurements of evoked activity indicate that synaptic physiology is unaffected by the electrodes, giving further support to the application of the proposed technology in the fabrication of advanced neurointerfaces. ACS Paragon Plus Environment

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Materials and Methods Electrode fabrication. Carbon fiber (CF)-based microelectrodes were fabricated as follows: individual 7 µmdiameter CFs (Goodfellow) were inserted into borosilicate capillaries (1.5 mm outer diameter, 0.86 inner diameter, A-M Systems) by applying a gentle suction. Capillaries were then pulled using a P-30 vertical puller (Sutter Instruments), so that the glass in the tip of one of the two resulting micropipettes was fused around the CF. Since CFs display a high tensile strength, they never got broken by pulling. The CF protruding from the fused tip of the micropipette was shortened to 1-2 cm using microscissors, and then cut to the desired final length (20, 50 or 250 µm) using a laser dissector microscope (Leica LMD6000). At the rear end of the micropipette, CF was trimmed at glass level. The first millimeters of the inner and outer surface of the rear end of the glass were painted with colloidal graphite (Agar Scientific) in order to provide adequate electrical connection with the CF. Electrodeposition of poly(3,4 ethylenedioxythiophene)/poly(styrenesulphonate)-co-maleic acid) (PEDOT-PSS-co-MA) on the surface of CF electrodes (hereafter referred to as CFEs) was performed by galvanostatic electropodeposition using a Bio-Logic SAS Versatile Single Potentiostat (VSP) in three-electrode cell configuration. CFE tips (working electrodes) were submerged into an aqueous potassium phosphate-buffered solution containing 15 mM 3, 4 ethylenedioxythiophene (EDOT monomer; SIGMA), and 20 mM poly (4-styrenesulfonic acid-co-maleic acid) sodium salt (PSS-co-MA; SIGMA) added as dopant compound. A platinum foil was used as counter-electrode, and a saturated calomel electrode (SCE) was used as reference electrode. A constant anodic current (1 µA/mm2) was applied for 960 seconds while monitoring the voltage at a sampling rate of 1 Hz. A polymerization charge of 96 mC/cm2 provides a polymer coating of 400-500 nm thickness on gold-coated glass slides20. Upon current application, potential in the working CFE quickly rose from basal value until reaching a plateau at approximately 0.9 V, and remained in such potential value until current switch-off. An example of EDOT:PSS-co-MA polymerization voltage curve obtained for a 250 µm-long electrode is shown in Supporting Information (S.I.) Fig. S1; similar results were obtained for 20 and 50-long ones.


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Polymer coating on the electrode surface was visualized by scanning electron microscopy (SEM), using a Philips XL 30 S FEG microscope in high vacuum mode and a Through Lens Detector at a working distance of 3.4-5.1 mm. Since CFs and PEDOT:PSS-co-MA are electroconductive, no further metallic coating was needed. The atomic composition of polymer coating was analyzed using energy dispersive X-ray spectroscopy (EDX, Oxford Instruments) integrated into a Nova NanoSEM200 microscope (FEI Company). Polymer electrical characterization. Success of electropolymerization and characterization of electrical properties of electrodes were also assessed by cyclic voltammetry (CV), a technique frequently used for the electrochemical study of neural electrodes21. CV was performed with the VSP in threeelectrode configuration, sweeping the voltage of the working electrode at 0.5, 5 and 50 V/s, from -1.05 to +0.55 V vs. a SCE reference electrode. The total electric charge (QCV) provided by electrodes was calculated as the absolute integral of the cathodic and anodic current enclosed in the cyclic voltammogram. Performance of electrodes under conditions of repetitive stimuli application was evaluated by chronoamperometry. For this, trains of 16 biphasic, 5 ms-width, square voltage pulses were administered through the electrodes at 0.2, 0.3, 0.4 and 0.5 volts while their current responses were recorded at 20 kHz sampling rate. Electrode biofunctionalization. Biofunctionalization was performed on CF-based, PEDOT:PSS-co-MA-coated electrodes (hereafter referred to as co-MAEs). Unlike PEDOT, PEDOT:PSS-co-MA exhibits carboxylic groups that allow the formation of covalent amide bonds with amino residues of proteins, thus making this doped polymer more appropriate for biofunctionalization. The molecule chosen for biofunctionalization was N-Cadherin (NCad), a calcium-dependent adhesion molecule that has a major role in axonal guidance and fasciculation in vivo22,23 and is required for neuronal migration24 and synaptic plasticity25,26. In cell culture, NCad induces profuse axonal extension and stimulates dendritic elongation and branching20,27-29. The procedure for NCad attachment to CP-coated carbon microfibers has been reported elsewhere20. In brief, a goat antibody raised against human IgG (Fc specific; Sigma) (antihuman IgG, AHIgG) was firstly bonded to co-MAEs. Carboxylic acid activating reagents N-hydroxysuccinimide

(NHS)

and

N-(3-dimethylaminopropyl)-N’-ethylcabodiimide


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(EDC) were employed to favor the formation of amide bonds between PSS-co-MA carboxylic groups and amino residues in the protein. Then, a recombinant human NCad-Fc chimera (R&D Systems) that combines human NCad with the Fc region of human immunoglobulin G1 (IgG1) was linked to the surface. Stability of biofunctionalization on PEDOT:PSS-co-MA- and NCad-coated electrodes (hereafter referred to as NCadEs) was verified by immunofluorescent detection of NCad on 250-µm (for ease of handling) electrodes after their use for recording of hippocampal spontaneous activity. The electrode tips were cut away using microscissors, and then fixed by one end to glass slides using tiny pieces of tape. Slides were incubated overnight with a specific monoclonal mouse anti-NCad antibody (Sigma) diluted 1:200 in phosphatebuffered saline (PBS) containing 0.5 % Triton X-100. After carefully washing with PBS (3 x 5 min), a polyclonal goat anti-mouse antibody conjugated with Alexa 488 (1:500, Fisher) was added for 2 hours. Slides were washed again (3 x 5 min) and coverslipped. Electrode recording properties characterization. CF-based electrodes (CFEs, co-MAEs and NCadEs) were used to record an externally applied voltage signal. An electrochemical cell specially designed to minimize charge dispersion between the electrodes was used for this purpose. Classic borosilicate micropipettes were pulled, filled with standard artificial cerebrospinal fluid (ACSF; composition: see below) (12-17 MΩ total resistance), and connected to the CV-7B headstage of a Multiclamp 700B amplifier (Axon Instruments) through a holder which uses a chlorided Ag wire to make electric contact with the pipette internal solution. Such micropipettes were used to apply square voltage pulses into the electrochemical cell filled with ACSF. The recording electrodes were placed at less than 100 µm from the tip of the stimulating micropipette. A first-cathodic (200 mV, 250 µs), then-anodic (100 mV, 500 µs) balanced-biphasic, slow-reversal pulse was designed using pClamp software. This waveform simulates the time course of extracellular spontaneous activity patterns obtained in models of rat motor cortex in vivo. Because the voltages and currents recorded through CFEs, co-MAEs and NCadEs had low magnitude (tens of microvolts and nanoamperes), the external voltage pulse was administered repeatedly (500 times at 1 Hz) and the recordings were averaged to reduce noise and enhance signal recognition.


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Electrophysiology. Postnatal p26-p35 Wistar rats were anesthetized with sodium pentobarbital and decapitated. The brain was quickly removed from the skull and dissected in cold ACSF. Horizontal brain slices (400 µm) were cut in cold ACSF using a vibratome (Leica), and then placed in a submerged holding chamber for at least 1 hour at room temperature (RT). The composition of the standard ACSF was (in mM): NaCl 120, KCl 2.5, NaH2PO4 1.0, MgCl2 1.2, CaCl2 2.5, NaHCO3 26.2, glucose 11, pH 7.4 when equilibrated with 95% O2 / 5% CO2. For recording, individual slices were transferred to a submerged chamber where they were superfused with ACSF at 28-29º C (for evoked activity recordings) or with modified ACSF (see below) at 36-37 ºC (for spontaneous activity recordings) at a constant rate (2-3 ml/min). ACSF temperature in the bath was adjusted with an inline heater (Warner Instruments). Spontaneous activity was recorded in stratum pyramidale of CA3 area. Low extracellular Mg2+ concentration increases the incidence of epileptiform activity in brain slices including the hippocampus30; hence, a modified, Mg-free ACSF was used to favor the incidence of sharp wave (SWR) activity. Its composition was as follows (in mM): NaCl 129, KCl 5, NaH2PO4 1.25, CaCl2 1.6, NaHCO3 21, glucose 10. Evoked field excitatory postsynaptic potentials (fEPSPs) were recorded in stratum radiatum of area CA1, where the Schaffer collateral/commissural afferents arising from CA3 pyramidal cells make synaptic contacts with the apical dendritic trees of CA1 pyramidal cells31. Stimuli (0.1 ms pulse width) were delivered to the Schaffer afferents at 0.1 Hz through a concentric bipolar tungsten-parylene electrode (A-M Systems, Inc) using a constant current isolated stimulator (Cibertec). Delivered currents ranged from 0.1 to 1.0 mA for input/output experiments. For analysis of synaptic facilitation, pairs of pulses were administered at eight interstimulus intervals, ranging from 10 to 500 ms; paired-pulse facilitation was calculated as the ratio of the second fEPSP slope divided by the first fEPSP slope. CF-based electrodes were connected through a holder coupled to the input headstage CV7B (Axon Instruments) of the Multiclamp 700B amplifier. The standard headstage-coupled holder was modified so that the Ag/AgCl wire could make appropriate contact with the conducting colloidal graphite painting at the rear end of the CF-based electrodes. Signals were amplified and sampled with a Digidata 1440A converter (Axon Instruments).


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Sampling frequency was 10 kHz for evoked potential recordings, 20 kHz for spontaneous activity recordings, and 25 kHz (maximum allowed system capacity) for recording experiments without tissue. Data were collected using pClamp10 software (Molecular Devices). Amplitude of spontaneous multi-unit activity (MUA) and SWR activity was calculated as the maximum peak-to-peak amplitude in the regions of interest of the recordings. Since SWR burst amplitude was always larger than MUA, SWR amplitude was calculated along the whole length of the recording. MUA amplitude was calculated in the time segments between SWR bursts. Statistical methods. All reported values are presented as average ± standard error of the mean. Statistical analysis between groups was performed by one-way ANOVA. p-values under 0.05 are considered statistically significant; higher levels of significance (

Carbon Microfiber Electrodes

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