List of groups and companies that have developed a particular MEA technology News on technological developments may be found on the continuously updated publication list on Multi Channel Systems’ website.
1. Introduction
Neural prostheses are devices that interface with the central or peripheral nervous system. They target at the capture, modulation or elicitation of neural activity, in most cases to record the information flow within a neural pathway for its online or later decoding, or to mimic or replace neural functionality that has been compromised or lost. While in theory any information carrying modality of the neuron could be tapped into ( Theoretically, any physical variable correlated to the membrane potential may be measured or altered. In a static scenario (e.g., resting potential) it could be the electrical field, in a dynamic scenario (e.g., upon de- or repolarization) its fluctuations, the (ionic) current(s) or even changes in the local magnetic field. Electrogenic cells in animals are neurons, muscle cells, pancreatic α- and -cells, kidney fibroblasts or electroplaques. Besides the light-induced electron separation process during photosynthesis in plants, some microorganisms and algae are capable of electrogenesis as well (Logan, 2009; Rabaey & Rozendal, 2010). The electrical field at an electrode site is usually quite distorted due to the non-homogeneity of the biological environment in its vicinity. This does not only complicate signal source analysis in neural recordings, but it also limits the spatial precision with which neurons can be stimulated electrically. Even worse, if an electrical stimulus triggers an action potential in an axon, it may spread in both directions (towards the synaptic arbor and the soma), which is not observed in natural neural activity propagation. bbreviations: A/D, analog-to-digital; AP, action potential; APS, active pixel sensor; ASIC, application specific integrated circuit; CMOS, complementary metal oxide semiconductor; CNT, carbon nanotube; CP, conducting polymer; CPFET, cell-potential field-effect transistor; CSC, charge storage capacity; CT, computer tomography; CV, cyclic voltammetry; D/A, digital-to-analog; DIV, days in vitro; DRIE, deep reactive ion etching; ECoG, electrocorticography; EEG, electroencephalography; EGEFET, extended gate electrode field-effect transistor; EMG, electromyography; EOSFET, electrolyte oxide semiconductor field-effect transistor; ERG, electroretinography; FET, field-effect transistor; GND, ground (electrode); HMDS, hexamethyldisilazane; ITO, indium tin oxide; ISFET, ion-sensitive field-effect transistor; LCP, liquid crystal polymer; LFP, local field potential; LIGA, Lithographie, Galvanoformung, Abformung; MEA, microelectrode array; MEG, magnetoencephalography; MOSFET, metal-oxide-semiconductor field-effect transistor; MRI, magnetic resonance imaging; MTM, metal transfer micromolding; NCAM, neural cell adhesion molecule; NGF, nerve growth factor; NW, nanowire; PDMS, poly(dimethylsiloxane); PEDOT, poly(3,4-ethylenedioxythiophene); PFOCTS, trichloro(1H,1H,2H,2H-perfluorooctyl)silane; PET, positron emission tomography; PI, polyimide, PMMA, poly(methyl methacrylate); PPX, poly(p-xylylenes); PPy, poly(pyrrole); PS, poly(styrene); PTFE, poly(tetrafluoroethylene); PU, poly(urethane); PVA, poly(vinyl alcohol); S/N, signal-to-noise ratio; SAM, self-assembling monolayer; TMS, transcranial magnetic stimulation; VLSI, very-large-scale integration.
This chapter will review recent design trends for microelectrode arrays (MEAs) with an emphasis on flexible polymer devices, which may be exploited for neuroprosthetics. A brief synopsis on the history of
1.1. In vitro microelectrode arrays
To better understand the events at the cell-electrode interface, a variety of MEAs for the The lipid double layer, which constitutes the cell membrane, can be considered a dielectric. The membrane thus acts as a capacitor that has no metal plates. Nevertheless, ions from the intra- and extracellular environments just accumulate at the (at physiological pH) negatively charged hydrophilic headgroups of the phospholipids at both sides of the membrane. If, as for most cells, the intra- and extracellular ionic compositions are different, a potential will build up across the membrane. As with any interface, the distribution of ions will very likely not be homogeneous but, to a first approximation, resemble a Helmholtz layer (Butt et al., 2003). Thus, any additional charges (such as those at the surface of a metal electrode), which create an electrical field gradient in the vicinity of the membrane, will lead to a reorganization of these two electric double layers. Due to the difference in their distances from the electrode, this reorganization will affect the intracellular membrane interface less strongly than the extracellular membrane interface.
‘Passive’ MEAs | # of electr. |
Device type and electrode materials | |
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30 | Two rows of 7 µm2 electroplated Pt on Au/Ni electrodes on glass, insulated by photoresist (Thomas, et al., 1972) | |
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36 R,S64 R,S | Ø 10 µm Au-coated ITO tracks on glass, insulated with a thermosetting polymer | |
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32 R,S | Two rows of sixteen 10 µm2 electrodeposited Pt electrodes on Au tracks on glass, insulated by SiO2 (Pine, 1980) | |
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25 R + 6 S | 15 x 15 µm2 electroplatinized, e-beam evaporated Au on Cr recording electrodes and 40 x 40 to 120 x 120 µm2 stimulation electrodes on a Ø 40 mm glass coverslip, insulated by photoresist. Ultrasonically welded Au wires Ø 76 µm connection to miniature connector | |
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32 R,S | Ø 10 (25) µm Au/Ti on glass electrodes insulated by polyimide | |
32 R,(S) | Ø 13 and 15 µm electroplatinized Au/Ti electrodes on perforated polyimide, insulated by polyimide | ||
58 R,S | Ø 25 µm Au/Ti electrodes on SU-8 above microfluidic PDMS channels on glass | ||
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64 R,S | 10 x 10 µm2 or Ø 25 - 30 µm electroplatinized Au on NiCr electrodes on 125 µm thick polyimide insulated by polyimide | |
& |
24, 36, 72 R,S flex54, 60 R,S32 R+12 S256 R,S | Ti (or Au/Cr) tracks with TiN- or Pt- coated electrodes (usually Ø 10 or 30 µm) on glass or Au or Ti tracks with Au or TiN electrodes on (perforated) polyimide insulated by Si3N4 or polyimide | |
60 R,S (3D) | TiN-coated Ø 30 µm, 10 – 50 µm high electrodeposited Au/Ti pillar electrodes on glass insulated by Si3N4 | ||
|
6 R |
10 x 10 µm2 Pt on Ta electrodes with electroplated Pt (Ø 35 µm) on a perforated Si/SiO2/Si3N4 substrate insulated by Si3N4 | |
34 R,(S)(3D) | 47 µm high, 15 µm exposed vapor-deposited Pt-tip on Ta electrodes on a porous (35%) Si substrate insulated by Si3N4 | ||
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30 R,S | 1.3 – 3.2 mm long, 15 µm wide Au/Cu on perforated polyimide (Upilex/Kapton) film | |
28 R,S | 50 x 100 µm2 electroplated Au on Cu/Ni on polyimide (Kapton®) with 5 perfusion holes | ||
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64 R,S | Electroplated Pt-black on 50 x 50 µm2 ITO tracks on glass insulated by a silicone photoresist | |
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64 R,S | 50 x 50 µm2 Au/Ni or 20 x 20 µm2, 50 x 50 µm2 or Ø 50 to 70 µm electrodeposited Pt-black electrodes on ITO tracks on a glass carrier insulated by polyimide or polyacrylamide | |
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60 R,S (3D) | Pt, Au or ITO tracks with Pt or Au electrodes on glass; spike-shaped electrodes are available; SU-8 insulator | |
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124 R | 24 x 5 + 4 electroplatinized Au/Ti electrodes (Ø < 10 µm) on fused silica wafer with SiO2/Si3N4/SiO2 insulation stack | |
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64 | 50 x 50 µm2 Au/Cr electrodes on glass, insulated by “spin-on-glass” or photopatternable silicone | |
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60 R,S | Ø 22 – 30 µm vapor-deposited Pt/Ti on Pyrex glass or Si, insulated by Si3N4; some are spatially partitioned by 5 interconnected clustering wells (Ø 3 mm in SU-8) | |
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39R+49S60 R64 R | ITO tracks with Ø 28 - 36 µm or 36 x 36 µm2 Au- or Pt-coated electrodes on glass in a tissue-“conformal” arrangement, insulated by Si3N4 or SU-8 | |
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36 R,S | Ø 75 or 100 µm electroless Au-plated Ni/Cu electrodes on polyimide (Kapton®) insulated by an acrylic adhesive and polyimide | |
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12 – 13 R (3D) | Laser scribed, electroplated Ni/Cu/Pt-black electrodes (Ø < 10 µm) on a SU-8 microtower (fluidic) structure (< 500 µm height) on perforated fused silica insulated by parylene | |
25 R (spike, 3D)+ 25 R (planar) | 300 – 500 µm high, Ø 50 µm Au/Cr spike-tip or Ø 50 µm Au planar electrodes through metal transfer micromolding (MTM) in PDMS on SU-8, PMMA, or PU carrier with parylene insulator either selectively laser- and globally RIE- (CHF3/O2 plasma) deinsulated at the electrode sites, or applied during “capping protection” of the electrode sites. Pt-black plating | ||
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64 S,R | Pt/TiW on < 10 µm high, µm Ø Ag or sub-µm Ø Si (electroporation) microneedles on Si insulated by Si3N4 | |
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64 R,(S) | Ø 10 – 100 µm Au/Ti electrodes on glass sputtered with IrOx with SiO2, Si3N4, SiO2 insulation sandwich | |
30 R,(S) | Ø 3-5 µm apertures above Au/Ti electrodes on Si/SiO2 with a SiO2, Si3N4 insulator stack | ||
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16 or 54 R,(S) + 2 GND | Drop cast CNT-decorated Ø 30 - 40 µm Pt/Ti recording and 2500 μm x 1000 μm GND electrodes (hexagonally arranged) on glass insulated by SiO2, Si3N4 | |
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62 R(3D) | Spine-shaped gold protrusions, electroplated on patterned Cu on glass, insulated by a SiC/Si3N4/SiO2 stack | |
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16 | 40 x 40 µm2 Au/Ti electrodes on glass and 20 x 20 µm2 SU-8 microwells and interconnecting micro-trenches; SU-8 insulation layer | |
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4 - 16 (R),S | Ø 250 µm Au-coated nail-head pins and liquid Ga/In (75.5/24.5) tracks in PDMS microchannels | |
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64 R,S + 2 S (6 768 R,S) |
Ø 30 µm Pt-black or Au/Ti electrodes on glass insulated by SU-8 or SiO2 | |
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52 R + 2 GND | Ø 35 µm Pt electrodes with interdigitated electrodes and PT1000 T sensor on glass insulated by Si3N4 | |
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20 R,S | 30-75 x 100 µm2 Au/Cr electrodes on deformable polyimide or PDMS with photo-patternable polyimide or PDMS insulator | |
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60 R,(S) | Ø 30 µm Ti electrodes on glass insulated by polystyrene (PS) on hexamethyldisilazane (HMDS) | |
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64 R | Electroplated Pt-black on Ø 30 µm Pt/Ti electrodes with Ø 5 or 10 µm substrate through-holes on backside-thinned Si/SiO2 carrier with microchannels insulated by parylene-C | |
‘Active’ MEAs | # of electr. |
Device type and electrode materials | |
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16 R64 R | 28 x 12 µm2 and 10 x 4 µm2 p-channel electrolyte oxide semiconductor FETs (EOSFETs) or Ø 30 - 60 µm extended gate electrode FETs (EGEFETs) insulated by Si3N4 | |
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128 R | 2 x 4 arrays of 16 Ø 10 µm Au/TiW/Al electrodes on CMOS IC, insulated by Si3N4 | |
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4 - 16,384 R | Electrolyte-oxide semiconductor field-effect transistors (EOSFETs) with Ø 4.5 µm charge-sensitive spots at a density of 16000/mm2 on silicon chips or multitransistor arrays (MTAs) based on metal-oxide-semiconductor field-effect transistor (MOSFET) technology | |
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≤ 150 R,(S) | Straight or kinked, oriented p- and/or n-type Ø 20 nm silicon nanowires (SiNW) spanning about 25 µm between Ni (source and drain) or Cr/Pd/Cr metal interconnects insulated by Si3N4 or PMMA | |
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4096 R | CMOS APS with 21 x 21 µm2 Al electrodes with optional Au-coating (electroless deposition) | |
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128 R,S | 8 x 16 array in CMOS technology with shifted Ø 10 - 40 (30) µm Pt-black (electrodeposited) on Pt (sputtered) electrodes insulated by an alternating Si3N4/SiO2 stack | |
126 R,S out of 11,011 | 128 x 128 array of Ø 7 µm shifted Pt (sputtered) electrodes at a density of 3150/mm2 on switch-matrix array in CMOS technology insulated by an alternating Si3N4/SiO2 stack | ||
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16 R | 4 x 4 recording sites each with 6 parallel silicon nanowire (SiNW) FETs with widths of 500 nm and pitch of 200 µm on metalized, doped Si source/drain contacts insulated by SiO2 | |
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100 R | 100 µm long silicon nanowires (SiNWs) with 30 x 40 nm2 rectangular cross section attached to Al on Si contact pads insulated by Si3N4 | |
24 R | CMOS with Au-coated Al electrodes | ||
‘Hybrid’ MEAs | |||
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32 R,S + 4 GND | Electrochemically platinized Ø 14 µm Au/Cr electrodes on Si/SiO2 carrier connected to a CMOS/VLSI amplifier chip, insulated by Si3N4 | |
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Multiparametric sensor with 6 µm2 CPFET and Ø 10 µm Pd or Pt electrodes, ISFET pH electrode, interdigitated impedance electrodes, photodiodes, oxygen and T sensor | ||
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61 R,S512 R,S519 R,S | Ø 2 - 5 µm electroplated Pt electrodes on ITO on glass, insulated by Si3N4, wire-bonded to ASIC readout & stimulation circuitry | |
61 R,S(3D) | Hexagonally arranged, ≤ 200 µm high, partially hollow W needles with electroplated Pt-tips, insulated by SiO2 and back-side connected to Al tracks, wire-bonded to ASIC readout & stimulation circuitry |
1.2. From in vitro to in vivo
From a conceptual point of view, the readout and stimulation physics of
Yet Apart from reaction products stemming from device degradation, exposed implant materials may also have catalytic properties. This statement does not exclude the temporary or permanent chemical and/or topographical device functionalization for manipulating or triggering a cell or tissue response in a controlled way (e.g., to support device – tissue integration or tissue regeneration).
Another critical issue is the actual interconnection of the devices to the outside world. Signals are commonly transferred to extracorporeal signal conditioning and processing electronics by cables. Not only are the connection points between cable and device a source of failure due to the detachment of the cable ends to the device by chemical degradation or mechanical forces. Once passing through the skin or skull, the pass-through hole has to be well sealed and stabilized to not let contaminants pass and become a site of chronic infection, and to not let the cables move and thereby exert mechanical stress onto the surrounding tissue. Recent trends therefore target at the transmission of signals through the skull by telemetric technologies, which, however, pose new challenges with respect to miniaturization, circuit protection against humidity, and energy transfer for powering the telemetric electronics. And it does not provide a solution for stabilizing the internal circuit-to-MEA connection, which may still be exposed to movement-related stress.
1.3. Electrode arrays for in vivo electrophysiology
For getting a general overview on The most well-known neuroprosthetic devices are cochlear implants, retinal implants, spinal cord stimulators, deep brain implants and bladder control implants.
1.3.1. Design and fabrication aspects
In the 80’s of the last century, the boom in microfabrication technologies opened the door for designing elaborate multi-microelectrode arrays with spatially distributed recording or stimulation sites. However, the choice of carrier, conductor and insulator materials depends on (and is thereby limited by) the often harsh fabrication and processing conditions. Given that materials in tissue-contact also need to fulfill the condition of being biocompatible (that is foremost, not being cytotoxic) and biostable (that is, not become degraded by the physiological environment), the number of suitable materials is quite low. Almost all of them are considerably more rigid than soft tissue, and the devices made from them tend to have sharp edges. Anyone ever having experienced a splinter in the thumb will remember how painful While the brain processes information on pain in other body parts, it does not feel pain itself.
The need for designing more flexible electrode arrays was already addressed in the 60’s of the last century (Rutledge & Duncan, 1966). Various strategies have been suggested since then to overcome some of the above mentioned limitations, particularly in the context of designing cochlear, retinal and deep brain implants. Today, the most commonly used flexible carrier and track insulation materials are polyimides (PIs), poly(p-xylylene) (PPX, and in particular poly(chloro-p-xylylene) (Parylene®-C)), poly(dimethylsiloxane) (PDMS), poly(tetrafluoroethylene) (PTFE), and occasionally less flexible liquid crystal polymers (LCPs) or photoresists ( Sometimes less noble metals like Ni or Cu are used as track and connection pad conductors. If the insulation layer of the probe has defects, they may partially dissolve into cytotoxic ionic species.
We therefore decided to deviate from common microelectrode array fabrication paradigms by resorting to a microchannel replication strategy with conductive polymers (CPs) completely replacing metals. This does not only allow the implementation of one and the same electrode layout in different types of insulating polymer backbones ( Alternatively, arbitrary 3D shapes could be inscribed into a photo-patternable polymer when resorting to UV laser or multi-photon lithography. One of the advantages would be the direct generation of non-vertical (e.g., conical) structures (Li & Fourkas, 2007; Thiel et al., 2010). If device geometries are simple and dimensions stay above a few tens of microns, even the master can be fabricated in the laboratory, as demonstrated by several groups (e.g., (Mensing et al., 2005)).
The bi-level master (1, grey, orange) ( To generate a Teflon®-like, fluorine-terminated anti-stick film, the wafer with the SU-8 microstructure or its epoxy copy can be either exposed for 5 minutes to C4F8 in a reactive ion etcher or to trichloro(1H,1H,2H,2H-perfluorooctyl)silane (PFOCTS) for one hour in a desiccator. One strategy to pull 50 - 200 µm thin PDMS microchannel scaffolds from their molding masters without tearing them apart was to place coated overhead or inkjet transparencies with their coated sides onto the uncured PDMS. In most cases, the water-soluble coating had a texture, which seemed to physically entrap the PDMS. After curing, the PDMS thus stuck to the sacrificial support transparency. It would detach from it automatically upon dissolution of the coating during immersion into water or ethanol for a few hours, leaving the texture topographies imprinted in the PDMS surface. As already briefly mentioned earlier, the transparency or its coating can be purposefully micro- or nanostructured to permanently transfer topographical cues into the PDMS and/or the CP electrodes.
During the
Both scenarios give useful devices for different application needs. A neuron may just grow its processes into a CP-coated channel without electrode membrane (Fig. 2, a3, left channel). Thus, the recorded signal can be easily attributed to that very neuron. Alternatively, if the well diameter were reduced to a few µm, a planar-patch-like recording array could be created. As recently reported by Klemic et al. (Klemic et al., 2005), Chen and Folch (Chen & Folch, 2006), Tonomura et al. (Tonomura et al., 2010), and Hofmann et al. (Hofmann et al., 2011), neurons tend to seal such microapertures to result in high signal-to-noise (S/N) ratios and selectivity. A back-filled channel with an intact electrode membrane (Fig. 2, a3, right channel) or a gPDMS composite (Fig. 2, b1-3) will act like a classical electrode instead. Even if the CP membrane were missing, the electrode would be defined by the nm-thick ring-shaped CP coating at the end of a channel. It is still an open question whether the impedance of these ring electrodes is sufficiently low to capture or induce neural signals.
1.3.2. PDMS as a soft and flexible substrate material
The platinum-catalyzed addition-crosslinking of vinyl-endblocked silicone polymers and silicone polymers with SiH functionality gives medical grade polydimethylsiloxane (PDMS) ( PDMS adheres strongly to itself. For a successful PDMS replication from a PDMS master, the master surface needs to be coated with an anti-stick layer. The one-hour exposure to trichloro(1H,1H,2H,2H-perfluorooctyl)silane (PFOCTS) in a desiccator results in a reusable Teflon®-like transferring layer (Zhang et al., 2010).
1.3.3. Flexible, polymer-based electrode materials
The desire of incorporating electronics into bendable and stretchable lightweight consumer devices (
While graphite and, in particular, carbon black are usually not categorized as polymers, they share some of their properties with respect to their extended carbon backbone. For their high electrical conductivity, biological inertness, low price and easy handling, they are excellent filler materials for creating flexible, voluminous conductor tracks or coatings with silicones or polyurethanes as the matrix (Calixto et al., 2007; Huang et al., 2011). As with any conductive filler ( Percolation as a mathematical concept refers to the long-range connectivity and its nature in random systems. The percolation threshold is the critical value of the (volume) occupation probability where infinite connectivity, in this case between conductive particles, first occurs. The resistance of thin-film electrodes may considerably deviate by two to three orders of magnitude from that of the bulk conductor material (Hu et al., 2006). Nevertheless, any wire deformation will alter the resistance of a wire. This phenomenon is exploited in strain gauge sensors. However, while the working principle of a strain gauge sensor relies on the mechanically induced changes in the cross-section geometry of the conductor, the resistivity of a composite material such as gPDMS seems to be dominated by the number of parallel conductive pathways. While the resistance in a strain gauge sensor increases with strain, the resistance of carbon- or silver-blended PDMS was actually found to decrease upon stretching for the better contact of the conductive particles (Niu et al., 2007).
1.3.4. Electrode functionalization and post-processing strategies
The postprocessing of devices serves two main goals:
the improvement of the electrical characteristics of the electrodes (mainly with respect to the decrease of their electrical impedances and/or the increase of their reversible charge delivery capacity (CDC)
Often, the reversible CDC is also referred to as the reversible charge storage capacity (CSC) or the save/reversible/capacitive charge injection limit (CIL).
the enhancement of their biocompatibility for their better tissue integration.
The electrical impedance is a more general concept of electrical resistance; it describes the frequency-dependent resistance of an electrical conductor. At ‘0 Hz’, the alternating current (AC) impedance of an electrode is identical to its direct current (DC) resistance. Over the physiologically relevant frequency range between 0.5 - 100 Hz (relevant for slow oscillations as in local field potentials (LFPs)) and 1 - 5 kHz (for the capture of individual action potentials (APs) of neurons The reasoning is as follows: During the firing of an action potential, the depolarization of the cell membrane lasts for about 1-2 ms. The temporal width of the extracellularly recorded component of such action potential is usually 1 ms (or less). This translates into a theoretical frequency of 1 kHz (or above) because 1000 (or more) such components will fit into 1 s. In a first approximation, the electrode can be considered the plate of a parallel-plate capacitor. Its electrical capacitance C is then directly proportional to the real electrode surface area A, which is equal or bigger than the geometrical electrode area. (C = ∙A/d, with the permittivity of the dielectric and the distance d between the plates). While the proportionality holds, the situation at the electrode-cell interface is certainly more complex: The second plate is not of the same type as the electrode but the cell membrane with a different surface area. Furthermore, the dielectric, the medium between electrode and cell membrane, is not static, and thus its permittivity is not a constant. Two types of currents are distinguished: Capacitive currents just charge the electrode; electrons will accumulate on the outer electrode surface without being injected into the electrolyte. Most stimulation electrodes are designed to be capacitive. In contrast, Faradaic currents pass the electrode-electrolyte interface. Because free electrons cannot be dissolved in aqueous environments, they become immediately involved in a redox-reaction. The occurrence of such undesirable reactions leads to chemical products that alter the composition of the physiological environment. To avoid any Faradaic currents, stimulation electrodes can be sealed by dielectric films (e.g., TiO2, Ta2O5 and BaTiO3). They are generated by oxidizing the respective metal electrodes, sputtering, sol-gel deposition or precipitation from organic or water-based dispersions. For an in-depth discussion see (Merrill et al., 2005; Cogan, 2008; Zhou & Greenberg, 2009; Merrill, 2010).
Many works have addressed the issue of providing signaling cues on electrode and substrate surfaces to mask the non-biological properties of a material, and to alleviate the acute and chronic disturbances imposed by a neuroprosthetic device onto its surrounding biological environment (Leach et al., 2010). By adsorbing polycations onto ready-made CP layers (Collazos-Castro et al., 2010) or by entrapping or intermingling cell adhesion and differentiation promoting proteins or their fragments (
1.3.5. Performance of polyMEA devices
The electrical, mechanical and optical characteristics as well as the recording performance of
Judging from the low phase of the impedance at low frequencies, the electrodes of the
We tested 80 µm PEDOT:PSS and gPDMS electrodes on dense carpets of cortical and hippocampal neurons from rats in cultured neural networks for their suitability as voltage-controlled stimulation electrodes without success Stimulation was based on charge-balanced biphasic voltage pulses not exceeding ±900 mV and 100 µs duration. Both polarity sequences (+, then -; -, then +) were tried. For a circular section of the cell membrane with a diameter matching that of commonly used electrodes (10 - 50 µm), the separated charge Q on the intracellular and extracellular side of the membrane during the resting (~ - 60 mV) or the action potential (~ 100 mV) is on the order of the CSC (several mC/cm2) of above mentioned stimulation electrodes.
1.3.6. Interconnection technologies
Connecting a microelectrode array to any kind of electronics is a critical issue. The mechanical clamping of contact pads requires a mechanism that is difficult to miniaturize and which might simply detach. Classical (wire) bonding introduces materials with limited flexibility that, due to stress and/or the contact chemistry between the different materials (including the humidity absorbed by a packaging compound), become the location of corrosion and break. By design, Non-uniform, wedge-like shaping of a device only requires the covering of the non-cured PDMS backside insulation by a sheet positioned in a ramp-like configuration during its curing. For acute ramp angles, adhesion forces between the PDMS and the sheet will prevent PDMS efflux.
1.3.7. Shielding
Because long, and in particular, high-impedance wires may act like antennas which tend to pick up noise from the environment, they are usually avoided. Instead, high-to-low impedance conversion electronics are placed as closely to the electrodes as possible (as in ‘active’ MEA devices). However, today, the rigidity of any type of conversion electronics would still compromise device flexibility. Therefore, proper shielding (like in coaxial cables) remains the only alternative. Also in this case, PEDOT:PSS or gPDMS may substitute graphite-based conductive lacquers to create a mechanically more flexible, tightly device-wrapping shielding. When graphite is mixed into non-cured PDMS, the viscosity of the paste, once it has reached the desired conductivity, can be decreased temporarily with solvents such as iso-propanol. The external surface After a brief exposure of PDMS to oxygen plasma, PDMS can be permanently bonded to itself. In this case, a short oxygen-plasma treatment of the device will enhance the adhesion of the shielding layer to the PDMS surface. Such plasma exposure did not have any detrimental effect on PEDOT:PSS electrodes.
1.3.8. Other observations
For some not-yet understood and investigated reasons, PDMS seems to provide a more favorable surface for cell and tissue adhesion than other common culture substrates. In two acute slice experiments (retinal whole mounts) it was observed that, after an initial weigh-down by a nylon-stocking ensheathed platinum U-wire for enhancing the tissue electrode contact, the weight could be removed after 30 minutes without compromising the signal quality. While the PDMS and the electrodes had been coated with poly-D-lysine and laminin in a standard procedure for enhancing cell adhesion on MEAs, this type of stickiness had never been observed in our lab with insulation layers made of silicon nitride, silicon dioxide (glass) or polymer (photo-) resists such as SU-8 or polyimide. A similar observation has been reported by Guo Plasma surface activation parameters for all types of MEAs stayed in the following ranges: One to three minutes at 30 – 60 W at 2.45 GHz in a 0.2 – 0.4 mbar pure oxygen atmosphere.
Preliminary results indicate that PEDOT:PSS can be embedded into a polymer matrix made of polyvinyl alcohol (PVA), glycerin and a di- or tricarboxylic acid as a crosslinker to render the PVA insoluble. However, this composite of high transparency and largely uncompromised conductivity will slightly swell in an aqueous environment. While a change in device geometries within the body is generally undesirable, a slight swelling of a polymer and/or its (hydrogel) electrodes may actually be favorable to enhance the electrode-tissue contact after device insertion. The water-uptake of PDMS itself is very low (below 1%). However, this might be just sufficient to stabilize device position within the tissue.
1.3.9. Open issues
As mentioned before, PDMS is very hydrophobic. Therefore, it is not wettable by aqueous or polar dispersions of organic conductors. Oxygen plasma treatment will render the PDMS surface hydrophilic by creating and exposing hydroxyl, carboxyl and peroxide groups on its surface, though. Depending on the storage conditions, this hydrophilicity is temporally more or less stable (Donzel et al., 2001). Under ambient conditions, it will usually degrade rapidly after the first 30 minutes. It can be anticipated that with shrinking channel feature sizes, the presented method of filling these channels (by coating the entire scaffold backside with the CP dispersion and then scraping it from the plateaus after the partial evaporation of the solvent) may not necessarily work well anymore. However, by playing with the two extremes of wettability, a channel-only plasma treatment may solve the problem. By temporarily covering the
While the softness and flexibility of all-polymer MEAs is one of their main assets, they have one major drawback: a MEA will not be easy to insert into dense tissue. A removable insertion device may alleviate this problem, though. During device fabrication, stiffer insertion and guidance aids could be embedded into the polymer microchannel scaffold such as anti-stick-coated polymer or glass fibers, which would then be withdrawn once the
The used type of PDMS already breaks after 100% tear. Softer and more stretchable polyurethanes or silicones are available that would render the devices even more flexible and tear-resistant. However, most of them have still to be tested for their biocompatibility. And some of them are only milky translucent, which make them less suitable for concurrent cell imaging studies.
1.3.10. Optional strategies and future directions
The flexibility of PDMS can be exploited for fabricating spherically bent neural probes that may become useful as retinal implants or for electroretinogram recordings as described
Similar to classical soft-lithography, the microchannels could also serve for the assisted transfer of conductive patterns onto other carrier substrates (Fig. 6). When placed onto a (nano-porous) carrier (
PEDOT:PSS can be purchased as an inkjet-compatible formulation. Thus, the filling of the microchannels with the conductor could not only be further automated, but different thicknesses or blends be deposited in different regions.
Alternatively, after the local laser-assisted alteration of the PDMS channel surfaces and the autocatalytic deposition of a Pt priming layer (Dupas-Bruzek, Drean, et al., 2009), EDOT or other precursors could be polymerized electrochemically to give electroconductive electrodes and tracks. The microchannels and the PDMS itself could be furthermore exploited in controlled drug-release strategies (Fig. 7b) (Colas, 2001; Musick et al., 2009). Various neural implant design studies with included microfluidics have already been reported (
With or without taking advantage of microfluidic connectors, neural processes could grow into the microchannels (Fig. 7b, left) as demonstrated by various groups (Morin et al., 2005; Claverol-Tinture et al., 2007; Benmerah et al., 2009; Lacour et al., 2010). This would increase the likelihood of identifying the actual origin of the bioelectrical signals. Alternatively, the preloading of empty or CP-coated microchannels with slow-release (electroconductive) hydrogels carrying diverse drugs could be pursued (Fig. 7a), thereby steering neural differentiation, regeneration and activity with growth or signaling factors, alleviating probe insertion damage by antibiotics and anti-inflammatory drugs, or attenuating the formation of glial scars by mitotic inhibitors (Peppas et al., 2006; Guiseppi-Elie, 2010). In that case, the electrical conductivity of the gel should be sufficiently high to warrant the coupling between the neuron and the conductive PEDOT:PSS film covering the microchannel walls. Alternatively or in combination, the PDMS itself could be loaded with drugs that are either soluble in PDMS or stored in porous cavities in a local silicone co-formulation. Delivering organic drugs through polymeric microchannels bears the risk of undesirable dissolution and accumulation of the compounds within the polymer over time, though. PDMS is particular prone to absorb
2. Conclusions
Neuroprosthetic devices should mimic as best as possible the tissue they are placed into. The tissue would then accept them as its own or just ignore them. They should furthermore be chemically, mechanically and functionally time-invariant for uncompromised performance. Fabricating electrode arrays exclusively from soft polymeric materials may be one step into that direction. The innovative concept of filling bi- or multi-level microchannel electrode array scaffolds with polymer conductors opens several new routes for designing and fabricating neuroprosthetic devices not only on the laboratory bench, but also through existing replica mass production schemes (
Acknowledgments
Several talented students and researchers contributed to the development and evolution of the various polyMEA devices. Their enthusiasm, help and dedication is greatly appreciated. In particular, I would like to thank Angelika Murr, Sandra Wolff, Christian Dautermann, Stefan Trellenkamp, Jens Wüsten and Mario Cerino for their support in exploring diverse device designs including their fabrication and testing, Tanja Neumann, Simone Riedel, Marina Nanni, Francesca Succol, and Maria Teresa Tedesco for their assistance in cell culture preparation and maintenance, Evelyne Sernagor for the preparation of retinal whole mounts, Paolo Medini and Giuliano Iurilli for epicortical recordings, Laura Gasparini and Francesco Difato for support in microscopy, Giacomo Pruzzo for his advice and assistance in developing various electronic gadgets, Tommaso Fellin for insightful discussions on future polyMEA exploitation pathways, Christiane Ziegler for providing the startup infrastructure and research environment, Sergio Martinoia and the Dept. of Biophysical and Electronic Engineering at the University of Genoa for temporary lab access, and Fabio Benfenati for subsequent support in shaping the polyMEA research line. A generous PDMS sample provided by Wacker Chemie AG is highly appreciated. This work is supported by the Italian Institute of Technology Foundation.
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Notes
- Theoretically, any physical variable correlated to the membrane potential may be measured or altered. In a static scenario (e.g., resting potential) it could be the electrical field, in a dynamic scenario (e.g., upon de- or repolarization) its fluctuations, the (ionic) current(s) or even changes in the local magnetic field.
- Electrogenic cells in animals are neurons, muscle cells, pancreatic α- and -cells, kidney fibroblasts or electroplaques. Besides the light-induced electron separation process during photosynthesis in plants, some microorganisms and algae are capable of electrogenesis as well (Logan, 2009; Rabaey & Rozendal, 2010).
- The electrical field at an electrode site is usually quite distorted due to the non-homogeneity of the biological environment in its vicinity. This does not only complicate signal source analysis in neural recordings, but it also limits the spatial precision with which neurons can be stimulated electrically. Even worse, if an electrical stimulus triggers an action potential in an axon, it may spread in both directions (towards the synaptic arbor and the soma), which is not observed in natural neural activity propagation.
- bbreviations: A/D, analog-to-digital; AP, action potential; APS, active pixel sensor; ASIC, application specific integrated circuit; CMOS, complementary metal oxide semiconductor; CNT, carbon nanotube; CP, conducting polymer; CPFET, cell-potential field-effect transistor; CSC, charge storage capacity; CT, computer tomography; CV, cyclic voltammetry; D/A, digital-to-analog; DIV, days in vitro; DRIE, deep reactive ion etching; ECoG, electrocorticography; EEG, electroencephalography; EGEFET, extended gate electrode field-effect transistor; EMG, electromyography; EOSFET, electrolyte oxide semiconductor field-effect transistor; ERG, electroretinography; FET, field-effect transistor; GND, ground (electrode); HMDS, hexamethyldisilazane; ITO, indium tin oxide; ISFET, ion-sensitive field-effect transistor; LCP, liquid crystal polymer; LFP, local field potential; LIGA, Lithographie, Galvanoformung, Abformung; MEA, microelectrode array; MEG, magnetoencephalography; MOSFET, metal-oxide-semiconductor field-effect transistor; MRI, magnetic resonance imaging; MTM, metal transfer micromolding; NCAM, neural cell adhesion molecule; NGF, nerve growth factor; NW, nanowire; PDMS, poly(dimethylsiloxane); PEDOT, poly(3,4-ethylenedioxythiophene); PFOCTS, trichloro(1H,1H,2H,2H-perfluorooctyl)silane; PET, positron emission tomography; PI, polyimide, PMMA, poly(methyl methacrylate); PPX, poly(p-xylylenes); PPy, poly(pyrrole); PS, poly(styrene); PTFE, poly(tetrafluoroethylene); PU, poly(urethane); PVA, poly(vinyl alcohol); S/N, signal-to-noise ratio; SAM, self-assembling monolayer; TMS, transcranial magnetic stimulation; VLSI, very-large-scale integration.
- The lipid double layer, which constitutes the cell membrane, can be considered a dielectric. The membrane thus acts as a capacitor that has no metal plates. Nevertheless, ions from the intra- and extracellular environments just accumulate at the (at physiological pH) negatively charged hydrophilic headgroups of the phospholipids at both sides of the membrane. If, as for most cells, the intra- and extracellular ionic compositions are different, a potential will build up across the membrane. As with any interface, the distribution of ions will very likely not be homogeneous but, to a first approximation, resemble a Helmholtz layer (Butt et al., 2003). Thus, any additional charges (such as those at the surface of a metal electrode), which create an electrical field gradient in the vicinity of the membrane, will lead to a reorganization of these two electric double layers. Due to the difference in their distances from the electrode, this reorganization will affect the intracellular membrane interface less strongly than the extracellular membrane interface.
- Apart from reaction products stemming from device degradation, exposed implant materials may also have catalytic properties.
- This statement does not exclude the temporary or permanent chemical and/or topographical device functionalization for manipulating or triggering a cell or tissue response in a controlled way (e.g., to support device – tissue integration or tissue regeneration).
- The most well-known neuroprosthetic devices are cochlear implants, retinal implants, spinal cord stimulators, deep brain implants and bladder control implants.
- While the brain processes information on pain in other body parts, it does not feel pain itself.
- Sometimes less noble metals like Ni or Cu are used as track and connection pad conductors. If the insulation layer of the probe has defects, they may partially dissolve into cytotoxic ionic species.
- Alternatively, arbitrary 3D shapes could be inscribed into a photo-patternable polymer when resorting to UV laser or multi-photon lithography. One of the advantages would be the direct generation of non-vertical (e.g., conical) structures (Li & Fourkas, 2007; Thiel et al., 2010).
- If device geometries are simple and dimensions stay above a few tens of microns, even the master can be fabricated in the laboratory, as demonstrated by several groups (e.g., (Mensing et al., 2005)).
- To generate a Teflon®-like, fluorine-terminated anti-stick film, the wafer with the SU-8 microstructure or its epoxy copy can be either exposed for 5 minutes to C4F8 in a reactive ion etcher or to trichloro(1H,1H,2H,2H-perfluorooctyl)silane (PFOCTS) for one hour in a desiccator.
- One strategy to pull 50 - 200 µm thin PDMS microchannel scaffolds from their molding masters without tearing them apart was to place coated overhead or inkjet transparencies with their coated sides onto the uncured PDMS. In most cases, the water-soluble coating had a texture, which seemed to physically entrap the PDMS. After curing, the PDMS thus stuck to the sacrificial support transparency. It would detach from it automatically upon dissolution of the coating during immersion into water or ethanol for a few hours, leaving the texture topographies imprinted in the PDMS surface. As already briefly mentioned earlier, the transparency or its coating can be purposefully micro- or nanostructured to permanently transfer topographical cues into the PDMS and/or the CP electrodes.
- PDMS adheres strongly to itself. For a successful PDMS replication from a PDMS master, the master surface needs to be coated with an anti-stick layer. The one-hour exposure to trichloro(1H,1H,2H,2H-perfluorooctyl)silane (PFOCTS) in a desiccator results in a reusable Teflon®-like transferring layer (Zhang et al., 2010).
- Percolation as a mathematical concept refers to the long-range connectivity and its nature in random systems. The percolation threshold is the critical value of the (volume) occupation probability where infinite connectivity, in this case between conductive particles, first occurs.
- The resistance of thin-film electrodes may considerably deviate by two to three orders of magnitude from that of the bulk conductor material (Hu et al., 2006).
- Nevertheless, any wire deformation will alter the resistance of a wire. This phenomenon is exploited in strain gauge sensors. However, while the working principle of a strain gauge sensor relies on the mechanically induced changes in the cross-section geometry of the conductor, the resistivity of a composite material such as gPDMS seems to be dominated by the number of parallel conductive pathways. While the resistance in a strain gauge sensor increases with strain, the resistance of carbon- or silver-blended PDMS was actually found to decrease upon stretching for the better contact of the conductive particles (Niu et al., 2007).
- Often, the reversible CDC is also referred to as the reversible charge storage capacity (CSC) or the save/reversible/capacitive charge injection limit (CIL).
- The reasoning is as follows: During the firing of an action potential, the depolarization of the cell membrane lasts for about 1-2 ms. The temporal width of the extracellularly recorded component of such action potential is usually 1 ms (or less). This translates into a theoretical frequency of 1 kHz (or above) because 1000 (or more) such components will fit into 1 s.
- In a first approximation, the electrode can be considered the plate of a parallel-plate capacitor. Its electrical capacitance C is then directly proportional to the real electrode surface area A, which is equal or bigger than the geometrical electrode area. (C = ∙A/d, with the permittivity of the dielectric and the distance d between the plates). While the proportionality holds, the situation at the electrode-cell interface is certainly more complex: The second plate is not of the same type as the electrode but the cell membrane with a different surface area. Furthermore, the dielectric, the medium between electrode and cell membrane, is not static, and thus its permittivity is not a constant.
- Two types of currents are distinguished: Capacitive currents just charge the electrode; electrons will accumulate on the outer electrode surface without being injected into the electrolyte. Most stimulation electrodes are designed to be capacitive. In contrast, Faradaic currents pass the electrode-electrolyte interface. Because free electrons cannot be dissolved in aqueous environments, they become immediately involved in a redox-reaction. The occurrence of such undesirable reactions leads to chemical products that alter the composition of the physiological environment. To avoid any Faradaic currents, stimulation electrodes can be sealed by dielectric films (e.g., TiO2, Ta2O5 and BaTiO3). They are generated by oxidizing the respective metal electrodes, sputtering, sol-gel deposition or precipitation from organic or water-based dispersions. For an in-depth discussion see (Merrill et al., 2005; Cogan, 2008; Zhou & Greenberg, 2009; Merrill, 2010).
- Stimulation was based on charge-balanced biphasic voltage pulses not exceeding ±900 mV and 100 µs duration. Both polarity sequences (+, then -; -, then +) were tried.
- For a circular section of the cell membrane with a diameter matching that of commonly used electrodes (10 - 50 µm), the separated charge Q on the intracellular and extracellular side of the membrane during the resting (~ - 60 mV) or the action potential (~ 100 mV) is on the order of the CSC (several mC/cm2) of above mentioned stimulation electrodes.
- Non-uniform, wedge-like shaping of a device only requires the covering of the non-cured PDMS backside insulation by a sheet positioned in a ramp-like configuration during its curing. For acute ramp angles, adhesion forces between the PDMS and the sheet will prevent PDMS efflux.
- After a brief exposure of PDMS to oxygen plasma, PDMS can be permanently bonded to itself. In this case, a short oxygen-plasma treatment of the device will enhance the adhesion of the shielding layer to the PDMS surface. Such plasma exposure did not have any detrimental effect on PEDOT:PSS electrodes.
- Plasma surface activation parameters for all types of MEAs stayed in the following ranges: One to three minutes at 30 – 60 W at 2.45 GHz in a 0.2 – 0.4 mbar pure oxygen atmosphere.