Fully Printed Î¼-Needle Electrode Array from Conductive Polymer Ink

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Biological and Medical Applications of Materials and Interfaces

Fully Printed µ-Needle Electrode Array from Conductive Polymer Ink for Bioelectronic Applications Sabine Zips, Leroy Grob, Philipp Rinklin, Korkut Terkan, Nouran Yehia Adly, Lennart Jakob Konstantin Weiß, Dirk Mayer, and Bernhard Wolfrum ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.9b11774 • Publication Date (Web): 19 Aug 2019 Downloaded from pubs.acs.org on August 22, 2019

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ACS Applied Materials & Interfaces

Fully Printed µ-Needle Electrode Array from Conductive Polymer Ink for Bioelectronic Applications Sabine Zips,† Leroy Grob,† Philipp Rinklin,† Korkut Terkan,† Nouran Yehia Adly,† Lennart Jakob Konstantin Weiß,† Dirk Mayer,‡ and Bernhard Wolfrum*,†,‡

†

Neuroelectronics - Munich School of Bioengineering, Department of Electrical and

Computer Engineering, Technical University of Munich, Boltzmannstr. 11, 85748 Garching, Germany

‡

Institute of Bioelectronics (ICS-8), Forschungszentrum Jülich, 52425 Jülich, Germany

KEYWORDS: printed electronics, additive manufacturing, microelectrode arrays, conductive polymers, carbon nanotubes, bioelectronics, extracellular recording

ACS Paragon Plus Environment

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ABSTRACT: Microelectrode arrays (MEAs) are widely used platforms in bioelectronics to study electrogenic cells. In recent years, the processing of conductive polymers for the fabrication of three-dimensional electrode arrays has gained increasing interest for the development of novel sensor designs. Here, additive manufacturing techniques are promising tools for the production of MEAs with three-dimensional electrodes. In this work, a facile additive manufacturing process for the fabrication of MEAs that feature needle-like electrode tips – so called µ-needles – is presented. To this end, an aerosoljet compatible PEDOT:PSS and multi-walled carbon nanotube composite ink with a conductivity of 323±75 S m-1 is developed and used in a combined inkjet and aerosol-jet printing process to produce the µ-needle electrode features. The µ-needles are fabricated with a diameter of 10±2 µm and a height of 33±4 µm. They penetrate an inkjet-printed dielectric layer to a height of 12±3 µm. After successful printing, the electrochemical properties of the devices are assessed via cyclic voltammetry and impedance spectroscopy. The µ-needles show a capacitance of 242±70 nF at a scan rate of 5 mV s-1


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and an impedance of 128±22 kΩ at 1 kHz frequency. The stability of the µ-needle MEAs in aqueous electrolyte is demonstrated and the devices are used to record extracellular signals from cardiomyocyte-like HL-1 cells. This proof-of-principle experiment shows the µ-needle MEAs’ cell-culture compatibility and functional integrity to investigate electrophysiological signals from living cells.


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INTRODUCTION Microelectrode arrays (MEAs) are common tools in bioelectronic systems and have been used extensively for research on the electrophysiological behavior of cells.1–5 This includes, for example, the examination of cardiomyocytes and neurons on the network and single-cell level under electrical stimulation, as well as the influence of pharmaceutical or toxicological cues.6–10 Moreover, CMOS-based MEAs with a high spatial density of electrodes were shown to be powerful platforms to study electrogenic cells even down to the sub-cellular level.11–13 However, clean-room processing as a classical fabrication method has several disadvantages. The equipment is expensive and the production workflow is complex and time-consuming. Device layouts cannot easily be changed, the material choice is often limited, and the fabrication of three-dimensional features with high aspect ratios poses challenges. For example, the production of siliconbased Utah arrays used for in vivo recordings involves complicated etching processes for structuring the needle-shaped electrodes. As an alternative to classical fabrication methods, additive manufacturing technologies and direct writing techniques have been increasingly used for the production of


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bioelectronic devices. Inkjet printing and aerosol-jet printing are prominent examples for additive processing of functional inks and have been widely applied for the development of biosensing platforms and MEAs.14–24 In contrast to classical fabrication, additive manufacturing benefits from only depositing the desired material on the required userdefined layout, which can be easily modified. This saves resources when developing novel sensing platforms and reduces fabrication costs. Furthermore, these techniques can be used for three-dimensional patterning of functional inks in high aspect ratios and a short amount of time opening up new possibilities for MEA designs.20,25–27 In terms of materials, conductive polymers have gained increasing interest over the last years for the production of bioelectronic devices such as MEAs or organic field effect transistors.28–33 A prominent example is poly(3,4-ethylenedioxythiophene) poly(styrene sulfonate) (PEDOT:PSS), which features properties such as biocompatibility, moderate Young’s moduli, a large variety of strategies for functionalization with biomolecules, and uptake of water.24,34–36 Devices fabricated from PEDOT:PSS show a high volume-specific capacitance and low charge-transfer resistance in electrochemical measurements as the bulk material readily takes up ions from the surrounding electrolyte.37 The oxidation or


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reduction state of PEDOT:PSS modulates the uptake of ionic species form solution, which can govern biological processes such as cell growth.38,39 To exploit these positive aspects, PEDOT:PSS composites have been used as coatings on clean-room fabricated MEAs and microneedle-array-based dry electrodes.40–45 A photolithographic lift-off process was developed to integrate PEDOT:PSS with parylene C for the fabrication of conformable

electrode

arrays.29

Replica

molding

of

PEDOT:PSS

involving

photolithographically produced templates was applied to pattern flexible MEAs for in vitro and in vivo action potential recordings.46 More recently, novel bioelectronic sensing platforms that employ three-dimensional PEDOT:PSS electrodes were developed. For instance, PEDOT:PSS nanorods were molded from polydimethylsiloxane (PDMS) stamps and used to electrically trigger the capture and release of tumor cells.47 The selective swelling of drop-casted PEDOT:PSS on photolithographically patterend MEAs was exploited to fabricate soft micropillar electrodes for recording from cardiomyocyte-like cells.48 Here, additive manufacturing offers a straight-forward option to produce three-dimensional PEDOT:PSS structures with high aspect ratios. For example, direct writing techniques have been developed to print


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stretchable PEDOT:PSS interconnects and pillar arrays.27,49 In comparison to flat electrodes, three-dimensional PEDOT:PSS electrodes may provide a major advantage for detecting signals from more complex cell culture models such as layered tissue constructs or three-dimensional cell culture. When used in brain implants for in vivo recordings, MEAs may benefit from needle-shaped PEDOT:PSS electrodes since they are expected to exhibit less mechanical mismatch with the brain tissue for example when compared to rigid silicon-based microelectrode arrays. In this work, a facile additive manufacturing process of a MEA with three-dimensional needle-like electrode tips – so called µ-needles – is demonstrated. To this end, an aerosoljettable PEDOT:PSS ink with multi-walled carbon nanotubes (MWCNTs) was developed and then used in a combined inkjet and aerosol-jet printing process to fabricate the µneedle MEAs. First, conductive feedlines were inkjet-printed onto a flexible polymer substrate. After aerosol-jet printing of the µ-needle electrode tips with the PEDOT:PSSMWCNT ink, the dielectric passivation layer was inkjet-printed without the need of an additional alignment procedure to create the micrometer-sized electrode openings. After successful fabrication, the final devices were electrochemically characterized and then


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used in a proof-of-principle experiment to record signals from cardiomyocyte-like cells. This showed the functional integrity and cell-culture compatibility of the printed µ-needle MEAs.

RESULTS AND DISCUSSION PEDOT:PSS-MWCNT ink composition. The conductive polymer ink for aerosol-jet printing was prepared from dry PEDOT:PSS pellets and several additives in a multicomponent system of high and low boiling-point solvents. The fractions of each component were chosen to balance aerosol generation, coalescence of the aerosol droplets after deposition, fast drying, and stability of the three-dimensional needle-like electrode tips – so called µ-needles. Moreover, an adequate conductivity of the material and long-term durability under cell-culture conditions were required for the final application. To meet these prerequisites, the PEDOT:PSS-MWCNT ink was formulated on a solvent system of deionized water, glycerol, and ethylene glycol in a ratio of 8:1:1 (w/w/w). The co-solvents glycerol and ethylene glycol were used to enhance the ink’s conductivity.50


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The surfactant and adhesion promoter (3-glycidyloxypropyl)trimethoxysilane (GOPS) and the emulsifier carboxymethyl cellulose (CMC) were added to ensure the stability of the structures in aqueous electrolytes.35,51–53 The conductivity as well as the mechanical stability of the µ-needles was further improved by doping the ink with MWCNTs, which have previously been demonstrated to enhance electrical and mechanical properties of composite films.34,54–59 The final PEDOT:PSS-MWCNT ink was used to aerosol-jet print three-dimensional µneedles (see Figure 1a). A close-up SEM image of a printed film shows that the MWCNTs remained intact in the deposited material (compare Figure 1b). The conductivity of the ink was assessed in a 4-point measurement setup and evaluated to 323±75 S m-1 (see Supporting Information for further details). Fabrication process of the µ-needle MEAs. After successful development of our PEDOT:PSS-MWCNT ink, we fabricated µ-needle MEAs using a combination of inkjet and aerosol-jet printing. The overall fabrication of the devices involved three main steps, which are described in the following.


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First, the conductive feedlines were inkjet printed with a silver nanoparticle-based ink onto a flexible polyethylene naphthalate (PEN) substrate (compare Figure 2a and b). After sintering the silver tracks, the ends of the feedlines were electrochemically plated with gold to avoid cytotoxic effects during later application in cell culture. Next, µ-needles were aerosol-jet printed onto the electrode tips using the custom-made PEDOT:PSS-MWCNT composite ink (see Figure 2c and d). The printing system generated the aerosol pneumatically using a pre-humidified nitrogen gas stream. To print the µ-needles, the aerosol stream was focused onto the desired position for 7 pulses of 4 s with waiting steps of 4 s in between pulses resulting in a total printing time of 52 s per µ-needle. The average diameter and height of the µ-needles was measured to 10±2 µm and 33±4 µm, respectively (aspect ratio of 3.2±0.5, n = 62 electrodes; mean ± standard deviation). The fabrication of high-aspect ratio features with the presented aerosol-jet printing technique is limited by the control of the material outflow and the drying time of the composite ink. Using an aerosol-jet printing system with an ultrasonic atomizer and increased temperature of the substrate-holder will potentially enable the reliable printing of significantly higher PEDOT:PSS-MWCNT structures.26


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Finally, the MEAs were passivated by inkjet-printing a continuous, UV-curable, polyacrylate-based insulating layer (compare Figure 2e and f). The print parameters of both the µ-needles and the dielectric layer were adjusted to yield a dielectric film lower than the height of the needles. Thus, the µ-needles protruded through the liquid dielectric and were not covered. No additional iterative alignment procedure was required in this step. This is in contrast to other printing approaches, which usually depend on precise alignment to successfully fabricate micrometer-sized electrodes.21–23 By eliminating this step, our approach drastically simplifies the production process for the printed MEA while maintaining micrometer resolution for the electrodes. After UV curing of the dielectric ink, the µ-needles penetrated the passivation layer to a height of 12±3 µm (see Figure 3a and S2), which resulted in a geometric surface area exposed to the electrolyte of 476±111 µm² (needle tip approximated as cylinder). A finished device is shown in Figure 3b. Each chip contained 12 electrodes with a pitch of 400 and 250 µm in x- and y-direction, respectively. Electrochemical characterization. After device fabrication, the electrochemical properties of the µ-needle MEAs were evaluated using cyclic voltammetry (CV) and


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impedance spectroscopy. Cyclic voltammograms were recorded by sweeping the working electrode potential between 0 and 0.6 V (vs Ag/AgCl reference electrode) in PBS electrolyte with different scan rates from 5 to 200 mV s-1 (see Figure 4a). The traces indicate a capacitively-dominated behavior with a slight pseudocapacitive contribution. For higher scan rates (>5 mV s-1), deviations from an ideal double-layer capacitance resulted in an increasingly non-rectangular shape of the voltammograms. This is in line with previous studies, which demonstrated similar behavior for porous carbon materials and conductive polymers.27,54,58,60–61 The lag in charging and discharging of the capacitor may be attributed to a change of the system’s time constant at higher scan rates caused by an increase of the distributed electrolyte resistances in the pores of the electrode.58,62– 64

Thus, the charge acceptance or delivery diminished with increasing sweep rate and the

electrodes exhibited an Ohmic drop in their current response. To further evaluate this behavior, the capacitance C of the µ-needle electrodes was calculated for different scan rates s using the difference of the anodic and cathodic current (𝑖a and 𝑖c, respectively) with the relation


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𝐶=

𝑖a ― 𝑖c 2

∙

1 𝑠

(1)

at a working electrode potential of 0.3 V.62 The capacitance of the µ-needle electrodes for the lowest scan rate of 5 mV s-1 and the highest scan rate of 200 mV s-1 was 242±70 nF (n = 23 electrodes) and 197±43 nF (n = 23 electrodes), respectively, which indicates a slight decrease in capacitance for increasing scan rates (see Figure 4b). Overall, both the general behavior and the calculated capacitances compare well to values for PEDOT:PSS composites previously reported in literature.27,53,64 To further investigate the performance of the electrodes, impedance spectroscopy was carried out. A low-amplitude signal was applied with a frequency from 1 Hz to 1 MHz. The impedance and phase curves revealed a combination of several RC elements, which is a characteristic response for composite materials with internally distributed contact impedances (compare Figure 4c).58,65–67 In addition, the impedance resulting from the interface of the polymer matrix and the porous metal current collector may have also been contributing to the RC responses. On a total of 6 chips, 62 out of 72 electrodes showed a similar behavior in impedance spectroscopy (yield of 86%), while 10 electrodes showed


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a significantly higher impedance and were thus considered as not functional. In the future, a better control of the material outflow will potentially increase the yield of functional µneedle electrodes. This may be achieved either by reformulating the ink composition or using an ultrasonic aerosol-jet printing system.26 For testing the long-term stability under cell-culture conditions, the impedance of the electrodes at 1 kHz was measured after incubation in PBS at 37 °C for up to 28 days (compare Figure 4d). On day 1, the electrodes showed an almost purely resistive behavior at 1 kHz (phase angle of –6°±1 °) with an impedance of 128±22 kΩ (n = 23 electrodes). After 28 days, 20 of the initial 23 electrodes remained functional and the impedance slightly decreased to 84±26 kΩ (n = 20 electrodes). Potentially, this change in impedance was caused by insulating components in the ink, for example GOPS or CMC, being washed out over the course of the incubation. All in all, the functionality of the µ-needle MEAs was not compromised when exposed to cell-culture conditions for longer periods of time. Recording of cell signals. Following electrochemical characterization, the µ-needle MEAs were used in a proof-of-concept experiment for the extracellular detection of signals from cardiomyocyte-like HL-1 cells.68 This cell line exhibits pacemaker activity and


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conduction of action potentials when grown to a confluent layer. After culturing the cells on the MEAs (n = 4) for 6 days, confluency was reached and cell signals were recorded amperometrically from all chips with an in-house-built amplifier system. The amplifier features a 1 GΩ feedback resistor and is similar to a system reported elsewhere.3 The cells either showed spontaneous conduction of action potentials or were stimulated with norepinephrine. After recording for several minutes, the signals were stopped by adding sodium dodecyl sulfate (SDS). An example time trace is shown in Figure 5a and the average spike shape of the extracellular signals in the recorded trace is depicted in Figure 5b. The HL-1 cells exhibited periodic action potential conduction with a frequency of 0.3 Hz, which is in line with previous reported results for this cell type.22,23 A maximum peak-to-peak amplitude of 188±7 pA (mean ± standard deviation of amplitudes recorded within one trace) was observed. In comparison, the root-mean-square (RMS) noise level was 8 pA, which resulted in a signal-to-noise ratio (SNR) of approximately 24. The measured minimum peak-to-peak amplitude and RMS noise on an electrode was 60±2 pA and 6 pA, respectively (approximate SNR of 9). We attribute variations in the signal amplitudes to heterogeneities in the cell layer. The phase delay between the


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observed signals from three neighboring electrodes on one chip is demonstrated in Figure S3, which indicates a rather local recording of the cell signals at the individual electrodes. A DIC image of a confluent cell layer growing on a µ-needle MEA is shown in Figure 5c. This proof-of-concept experiment demonstrated the cell-culture compatibility of the printed µ-needle MEAs and their applicability for electrophysiological recordings from living cells. In principle, amperometric recordings as used in this application allow the detection of electrochemical as well as electrophysiological signals. For example, electrochemical recordings at appropriately biased electrodes can be performed for the detection of dopamine neurotransmitter release.4 Here, our interest was focused solely on the detection of signals generated by the cardiomyocyte-like cells. Thus, we chose a moderate bias potential at the electrodes (0 V vs. Ag/AgCl) to avoid unwanted interference by oxidation of redox-active molecules such as norepinephrine. The extracellular voltage signals are capacitively coupled at the electrode interface generating a transient current, which is converted to a voltage signal by the amplifier circuit. Thus, a reduction of the electrode impedance is expected to yield larger signals. However, in


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contrast

to

extracellular

voltage

measurements,

the

noise

in

amperometric

measurements scales inversely with the impedance at the interface.3 Thus, an increase of the electrode size is likely to also contribute to increasing noise levels. Future investigations should be directed at exploiting the simultaneous recording of electrophysiological and electrochemical signals. In particular, modified µ-needle arrays could be used to investigate action potential generation and neurotransmitter release in 3D-cell culture systems.

CONCLUSION The present study demonstrates a facile additive manufacturing process for the production of MEAs with three-dimensional µ-needle electrode tips. To this end, an aerosol-jet compatible PEDOT:PSS and MWCNT composite ink with a conductivity of 323±75 S m-1 was developed. Employing a combined inkjet and aerosol-jet printing process, this ink was used to successfully fabricate the µ-needle electrode openings without the need of an alignment step to create micrometer-sized electrode features. The µ-needles were printed with a diameter and height of 10±2 µm and 33±4 µm, respectively


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(aspect ratio of 3.2±0.5) and penetrated the passivation layer to a height of 12±3 µm. After fabrication, the electrochemical properties of the µ-needle MEAs were assessed and the long-term stability in aqueous electrolytes was tested. The capacitance of the µneedle electrodes at a scan rate of 5 mV s-1 was 242±70 nF and the impedance at 1 kHz was 128±22 kΩ. Finally, the devices were used to record signals from cardiomyocyte-like HL-1 cells in a proof-of-principle experiment. Successful recordings of the cell signals demonstrated the µ-needle MEAs’ functional integrity and long-term stability under cellculture conditions. In conclusion, the PEDOT:PSS-MWCNT composite and fabrication of the µ-needle MEAs presented herein will be a useful tool for the development of novel bioelectronic devices. The implementation of passivation layers with arbitrary thickness may offer possibilities to study the influence of soft or rigid interfaces on two-dimensional cell cultures, while maintaining a planar (or three-dimensional) substrate profile. Moreover, the µ-needle electrode tips may be exploited in the future for research on the electrophysiological behavior of cells in a more complex environment, such as recording


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or stimulation of cell signals from the inside of layered or three-dimensional tissue-culture constructs.

EXPERIMENTAL Chemicals and materials: The PEN substrates Teonex® Q65HA and Q83 (DuPont Teijin Films, Hopewell, USA) were obtained from Pütz Folien (Taunusstein, Germany). Silver nanoparticle ink (Orgacon SI-J20x, Agfa, Mortsel, Belgium) was obtained from Ceradrop (Limoges, France). UV curable insulator ink was purchased from Dycotec Materials (DM-IN-7003-I, Swindon, UK). PEDOT:PSS dry pellets, CMC, GOPS, glycerol (≥ 99 %), ethylene glycol (≥ 99 %), Claycomb medium, norepinephrine bitartrate, trypsin, ethylenediaminetetraacetic acid (EDTA), fibronectin, and gelatin were acquired from Sigma Aldrich (St. Louis, USA). Fetal bovine serum, penicillin, streptomycin, and Lglutamine was purchased from ThermoFisher Scientific (Waltham, USA). PDMS Sylgard 184 (10:1 w/w base:curing agent) was bought from Dow Corning (Wiesbaden, Germany). Surfactant Dynol™ 604 was obtained from Evonik (Essen, Germany). Carboxylated MWCNTs (≥95 %, OD 15 nm, length 10-50 µm) were bought from IoLiTec Ionic Liquids


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Technologies (Heilbronn, Germany). 2-Propanol (≥ 99.5 %), ethanol (≥ 99.5 %), and ascorbic acid were purchased from Carl Roth (Karlsruhe, Germany). Aqueous gold potassium cyanide was obtained from Enthone-OMI (PUR-A-GOLD™ 401B; ‘sHertogenbosch, Netherlands). Deionized water was taken from an Ultra Clear® purification system (Evoqua Water Technologies, Barsbüttel, Germany). Phosphatebuffered saline (PBS, 1×) was purchased from Biochrom (Berlin, Germany) and SDS stock solution (SDS, 1 M in water) was prepared using dry tablets from Sigma Aldrich (St. Louis, USA). PEDOT:PSS-MWCNT ink: 0.35 % (w/w) CMC was dissolved in a mixture of water, glycerol, and ethylene glycol (8:1:1, w/w/w) and stirred for 10 min. Next, 1.25 % (w/w) PEDOT:PSS was added. The mixture was stirred again and then homogenized with an ultrasonic horn for 1 min (127.5 µm amplitude, 6 mm KE76 probe, HD 2200.2, Bandelin, Berlin, Germany). Next, 0.25 % (v/v) GOPS, 0.02 % Dynol 604 (v/v; from 1:1 w/w stock solution in ethanol), and 0.5 % (w/w) MWCNTs were added and the ink was further sonicated for 10 min.69 During sonication, the ink was cooled in an ice bath to avoid


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overheating. Subsequently, the PEDOT:PSS-MWCNT dispersion was centrifuged (ROTANTA 460 R, 10 minutes at 100 rcf ) and the supernatant was used for printing. Conductivity of the composite ink: The conductivity of the PEDOT:PSS-MWCNT ink was measured by printing electrodes with contact pads onto a Q83 substrate using an aerosol-jet printing system from Optomec (Albuquerque, USA; compare Figure S1) integrated in an inkjet printer from Ceradrop (F-series, Limoges, France). Detailed print parameters, fabrication steps, and optical data of the structures are given in the Supporting Information. A potentiostat (VSP-300, BioLogic Science Instruments, Seyssinet-Pariset, France) was used to apply a voltage ramp (0 to 0.6 V with 5 mV s-1) to the electrodes and the current was recorded in a four-point configuration. The conductivity was calculated following Equation S1. Fabrication of the µ-needle electrode arrays: Conductive feedlines were printed with a silver nanoparticle ink (Orgacon SI-J20x, Agfa, Mortsel, Belgium) on a state-of-the-art inkjet-printing system (F-series, Ceradrop, Limoges, France) using disposable printing cartridges (DMC 1 pl, Fujifilm Dimatix, Santa Clara, USA) onto Teonex® Q65HA at 250 Hz and 1 kHz jetting frequency and a drop spacing of 20 µm. During the print, the


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nozzle plate and the substrate holder were kept at 50 °C. After drying at 60 °C overnight, the prints were sintered at 150 °C for 1 h. The electrode arrays were then electrochemically covered with gold from gold cyanide salt solution using a potentiostat (VSP-300, BioLogic Science Instruments, Seyssinet-Pariset, France) in a three-electrode configuration with a platinum mesh counter electrode and a Ag/AgCl reference electrode (3 M NaCl, RE-6, BASi, West Lafayette, USA) by running 6000 pulses of 20 ms at -1 V followed by 70 ms at 0.4 V and 10 ms at 0 V. As a next step, µ-needles were aerosol-jet printed (Optomec, Albuquerque, USA) onto the tips of the feedlines using the PEDOT:PSS-MWCNT ink. Our system uses a pneumatic atomizer and a pre-humidified nitrogen gas stream. The µ-needles were fabricated with a push flow of 15 sccm (standard cubic centimeter per minute), a focusing ratio of 1, and a 150 µm nozzle.70 During printing, the temperature of the substrate holder was set to 58 °C. Alignment was done manually using the printer’s camera system. The µ-needles were printed by focusing the aerosol stream onto one position for 7 pulses of 4 s with drying steps of 4 s in between pulses by repeated opening and closing of the shutter. In addition, µ-needles were aerosol-jet printed onto a gold-covered wafer for characterization with SEM (NVision 40,


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Zeiss, Oberkochen, Germany). After printing, the solvents of the ink were evaporated at 150 °C overnight. The chips were then plasma treated (O2, 0.5 mbar, 190 W, 30 s, Diener Femto, Diener electronic, Ebhausen, Germany) and passivated by inkjet-printing a continuous layer of a UV curable insulator ink (DM-IN-7003-I, Dycotec Materials, Swindon, UK) at 55 °C nozzle plate and substrate-holder temperature, 1 kHz jetting frequency, and 50 µm drop spacing. The passivation layer was UV cured with a dose of approximately 1 J cm-2. The height of the needles was measured before and after passivation using a laser scanning microscope (VK-X250, Keyence, Osaka, Japan) and evaluated with MultiFileAnalyzer (Keyence, Osaka, Japan). Glass rings were then glued onto the chips with PDMS which was cured at 120 °C for 1 h. Electrochemical characterization of the µ-needle electrode arrays: CV and impedance spectroscopy was carried out using a potentiostat (VSP-300, BioLogic Science Instruments, Seyssinet-Pariset, France) in a three-electrode configuration with a platinum-wire counter electrode and a Ag/AgCl reference electrode (3 M NaCl, RE-6, BASi, West Lafayette, USA) in PBS electrolyte. The CV curves were recorded by cycling the potential of the working electrode between 0 V and 0.6 V with different scan rates of


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5, 10, 20, 50, 100, and 200 mV s-1 for 3 to 10 cycles. Impedance spectroscopy was performed by applying a sine signal with an amplitude of 20 mV and varying frequencies from 1 Hz to 1 MHz. For testing the long-term stability under cell-culture conditions, impedance spectroscopy was carried out after the chips were stored in PBS at 37 °C and 5 % CO2 in a humidified incubator for up to 28 days. The data was recorded with EC-Lab (BioLogic Science Instruments) and analyzed with MATLAB® (MathWorks, Natick, USA). The capacitance of the electrodes for the different scan rates was calculated using Equation 1. HL-1 cell culture: Cardiomyocyte-like HL-1 cells were cultured in Claycomb medium supplemented with fetal bovine serum (10 %), penicillin/streptomycin (both 100 µg ml-1), norepinephrine (0.1 mM in ascorbic acid solution) and L-glutamine (2 mM). The cells were grown at 37 °C in 5 % CO2 atmosphere in a humidified incubator and passaged after reaching confluency and displaying mechanical contraction. The cells were detached via incubation in protease solution (0.05 % Trypsin-EDTA). Prior to cell seeding, the samples were plasma treated (O2, 0.8 mbar, 304 W, 5 min, Diener Femto, Diener electronic, Ebhausen, Germany) and sterilized by dipping into 2-propanol. After drying, the chips


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were incubated with a solution of fibronectin (5 µg ml-1) and gelatin (0.2 mg ml-1) for approximately 1 h at 37 °C. The protein solution was aspirated and the chips were rinsed with PBS. The cells were then seeded onto the samples for confluency after 1-6 days in

vitro (DIV1-DIV6). DIC images of the cell layer were taken with an Axiovert 40CFL microscope (Zeiss, Oberkochen, Germany). Recording of cell signals: The extracellular signals were observed amperometrically on DIV6 (n = 4 chips) with an in-house built amplification system capable of recording from a maximum of 64 channels with a sampling rate of 10 kHz. The amplifier features a 1 GΩ feedback resistor and is similar to a system reported elsewhere.[3] The setup was grounded and shielded with a Faraday cage. The measurements were performed with a Ag/AgCl reference electrode and AC coupled with an effective bandwidth of 1 mHz to 3.4 kHz. Medium exchange was performed 1 h prior to the experiments and 5 mM norepinephrine stock solution (dissolved in 33.6 mM ascorbic acid) was freshly prepared. If the HL-1 cells did not exhibit spontaneous signals, they were chemically stimulated during the measurement with norepinephrine (5 µl from the stock solution pipetted into 1 ml medium on the chip). Finally the cell signal was stopped by adding SDS (100 µl from


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1 M stock solution pipetted into 1 ml medium on the chip). The total measurement time was approximately 1–5 min. The signal analysis was performed in MATLAB® (MathWorks, Natick, USA). A low-order polynomial fit was subtracted from the data to remove low-frequency baseline fluctuations.


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Figure 1. Aerosol-jet printing of PEDOT:PSS-MWCNT composite ink. a) SEM image of a printed µ-needle. b) Close-up SEM image of a printed film showing the MWCNTs in the deposited ink.


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Figure 2. Fabrication steps of the µ-needle electrode arrays. a) Inkjet printing of conductive feedlines with a silver nanoparticle ink. b) Microscope image of a chip with the printed silver tracks. c) Aerosol-jet printing of needle-shaped electrode tips in a PEDOT:PSS composite ink doped with MWCNTs. d) Microscope image of a printed µneedle on an electrode tip. e) Inkjet-printing of the UV-curable insulating layer. f) Electrode tip after passivation.


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Figure 3. a) Overlay of the surface scans of an electrode tip before and after printing the insulating layer. The height and width of the µ-needle is approximately 35 µm and 10 µm, respectively. b) Image of an encapsulated chip.


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Figure 4. Electrochemical characterization of the printed µ-needle electrodes. a) Cyclic voltammograms of an example µ-needle electrode in PBS electrolyte with scan rates of 5, 10, 20, 50, 100, and 200 mV s-1 (potential on working electrode EWE vs. current I). b) Calculated average capacitance C of the µ-needle electrodes (n = 23) for the different scan rates. c) Impedance spectroscopy data of a µ-needle electrode. d) Long-term stability test. Impedance measured at 1 kHz after keeping the chips for up to 28 days


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under cell-culture conditions. The number of functional electrodes is given in the parentheses.


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Figure 5. Signal recording from HL-1 cells. a) Example time trace of cell signals detected with a µ-needle electrode showing periodic signals with a frequency of 0.3 Hz. b) Average spike shape of all recorded spikes in the trace (± standard deviation). c) DIC image of an HL-1 cell layer growing on a µ-needle MEA.


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ASSOCIATED CONTENT

Supporting Information

The Supporting Information is available free of charge on the ACS Publications website.

Conductivity measurement of the PEDOT:PSS-MWCNT ink with Figure S1, S2, and S3 (PDF)

AUTHOR INFORMATION

Corresponding Author *E-mail: [email protected]

Notes The authors declare no competing financial interest.

ACKNOWLEDGMENT

The authors acknowledge funding by the Bernstein Center for Computational Neuroscience Munich (grant number 01GQ1004A, BMBF) and the DFG (grant number


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DFG INST95/1331-1). S. Zips received funding from the Friedrich Naumann Foundation for Freedom (stipend number 8217/P622). Furthermore, the authors thank E. BrauweilerReuters and M. Golibrzuch for SEM imaging, H. Url for HL-1 cell culture, and S. Inal and D. Khodagholy for discussions on material composition.

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TABLE OF CONTENTS GRAPHIC


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Fully Printed Î¼-Needle Electrode Array from Conductive Polymer Ink

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