Biomedical Microdevices

, Volume 13, Issue 3, pp 441–451

Novel multi-sided, microelectrode arrays for implantable neural applications


    • Department of Electrical EngineeringUniversity of Michigan
  • Nick B. Langhals
    • Department of Electrical EngineeringUniversity of Michigan
  • David J. Anderson
    • Department of Electrical EngineeringUniversity of Michigan
  • Daryl R. Kipke
    • Department of Electrical EngineeringUniversity of Michigan

DOI: 10.1007/s10544-011-9512-z

Cite this article as:
Seymour, J.P., Langhals, N.B., Anderson, D.J. et al. Biomed Microdevices (2011) 13: 441. doi:10.1007/s10544-011-9512-z


A new parylene-based microfabrication process is presented for neural recording and drug delivery applications. We introduce a large design space for electrode placement and structural flexibility with a six mask process. By using chemical mechanical polishing, electrode sites may be created top-side, back-side, or on the edge of the device having three exposed sides. Added surface area was achieved on the exposed edge through electroplating. Poly(3,4-ethylenedioxythiophene) (PEDOT) modified edge electrodes having an 85-μm2 footprint resulted in an impedance of 200 kΩ at 1 kHz. Edge electrodes were able to successfully record single unit activity in acute animal studies. A finite element model of planar and edge electrodes relative to neuron position reveals that edge electrodes should be beneficial for increasing the volume of tissue being sampled in recording applications.


Neural recordingMicroelectrode arrayParyleneNeural prosthesesDrug deliveryChemical mechanical polishing

1 Introduction

High density neural recording devices have been and continue to be a critical tool for neuroscientists studying complex neural circuits (Buzsaki 2004; Riera et al. 2009). This technology also enables potentially life-transforming clinical devices for the treatment of paralysis (Donoghue et al. 2007) or refractory epilepsy (Stacey and Litt 2008). Next generation neural probes must enable stable, higher density recordings to advance neurotechnology (Cheung 2007; Kipke et al. 2008). This work offers a new polymer-based microfabrication technique that enables a larger design space for electrode placement, electrode density, and structural flexibility.

Parylene-C (poly-chloro-p-xylylene) has desirable material properties for use in a microfabricated implantable device. It has the highest certification level of biocompatibility, USP Class VI. This biocompatibility results from it being chemically inert and synthesized using a solvent-less polymerization process (Gorham 1966). Critically important to microfabrication, one can generate conformal, pinhole-free barrier films that are subsequently dry-etched using an oxygen plasma. For these reasons, parylene-C has long been used in neurotechnology, especially as a coating. It has been successful in coating microwires (Loeb et al. 1977) and micro-machined silicon arrays (Kim et al. 2008). Researchers have also microfabricated planar devices with a parylene-C substrate in a microfluidic device with recording capabilities (Takeuchi et al. 2005), retinal stimulation array (Li et al. 2005; Rodger and Tai 2005), and a flexible nerve cuff (Spence et al. 2007).

Unfortunately, a microfabricated parylene-based neural probe, as opposed to a parylene coating on free-standing conductors, cannot be assumed to provide low noise performance. Researchers have shown that water and salts will rapidly diffuse to the various interfacial boundaries (Yasuda et al. 2001; Seymour et al. 2009). Interfacial delamination due to poor wet adhesion of parylene on metal (Sharma and Yasuda 1982; Yasuda et al. 1996) and mechanical stress may create low impedance pathways resulting in cross-talk or shunt loss. More recently, researchers have shown that thin parylene films (50-300 nm) have regular pinholes, which can be used for drug diffusion control (Zeng et al. 2005; Pierstorff et al. 2008; Robinson et al. 2008), and must be of concern in electrical insulation applications. High-density parylene-C microfabricated structures have closely spaced conductors similar to conventional silicon-based neural probes (Vetter et al. 2004; Wise 2005; Kim et al. 2008). Unlike their inorganic counterparts, parylene-C devices are prone to poor adhesion at either the dielectric to dielectric interface or at the dielectric to metal interface (Sharma and Yasuda 1982; Yasuda et al. 2001; Seymour et al. 2009).

Research on alternative parylene materials have promising results. Complementary reactive parylene layers of poly(4-aminomethyl-p-xylylene-co-p-xylylene) and poly(4-formyl-p-xylylene-co-p-xylylene) were discovered and synthesized in the laboratory of Joerg Lahann and found to improve adhesion in microfluidic applications (Chen et al. 2008). In another study we evaluated these same materials, PPX-CH2NH2 and PPX-CHO, and found they improved wet metal adhesion and maintained higher electrical impedance relative to parylene-C alone (Seymour et al. 2009). These reactive parylene materials are used to interface the metal traces in this work, however, more research is necessary to understand their mechanical and biological properties over time and across temperatures. Fortunately, one advantage of this process is that alternative dielectric materials can be used instead of the reactive parylene if necessary. Compatible dielectric materials are limited primarily by the temperature of deposition. Researchers have reported successful combination of polymer and low-temperature inorganic dielectrics such as PECVD SiO2, Si3N4 (Weaver et al. 2002; Lewis et al. 2004; Wuu et al. 2005) or SiC (Chiang et al. 2003; Cogan et al. 2003) all can be deposited below 220°C.

One electrode design in particular that drove the development of a new process was the concept of a double-sided edge electrode on an ultra-thin substrate. In a previous histological study (Seymour and Kipke 2007), we hypothesized that if a structural feature size is smaller than a reactive cell body (<7 μm), the resulting encapsulation would be mitigated by the prevention of cellular spreading. We investigated this relationship between size and tissue reactivity using parylene structures and found the cellular density around the sub-cellular feature was less than half of that around the probe shank. Our interpretation of the biological response suggested that an optimal electrode design is a two or three-sided microelectrode on the edge of thin parylene substrate (5-μm thick). We present here the first microfabrication approach capable of implementing such a design which we hope may be used in future long-term performance studies.

In general, the scope of this work was to develop a polymer-based technology platform that enabled a greater design space for microelectrodes and structural features. The platform needed a dynamic structural range including a thick shank for insertion strength or reservoir capabilities, as well as thin, flexible structures where electrodes could be positioned. The electrode placement also needed to be top-side and bottom-side to enable the edge electrode and higher density arrays. Finding a planarization process that enabled the implementation of double-sided electrodes was the critical process step. We used chemical mechanical polishing (CMP) of parylene-C over SiO2 to construct a planar insulation layer with electrode openings facing toward the wafer. CMP of parylene is rare, but one investigation showed that it was possible with parylene-N and resulted in only a small increase in oxygen content (Yang et al. 1997). CMP of polyimide is more common and may be a useful alternative (Hyoung-Gyun et al. 2000). We found the most significant challenge in polishing parylene was the prevention of delamination, which we discuss in detail. Finally, we successfully recorded neural spike activity in the rat neocortex in acute experiments.

2 Experimental

2.1 Microfabrication

The microfabrication for these structures required six masks and was conducted on 4” silicon wafers in the Lurie Nanofabrication Facility at the University of Michigan. For simplicity, we describe the process in eight steps (Fig. 1(a-h)). A variety of neural devices were included in the mask set but the cross-section shown in the illustrated process was taken from a polytrode recording array that included face-down electrodes, face-up electrodes, and edge electrodes (Fig. 3(a)).

Figure 1(a). Five microns of high temperature (910°C) SiO2 was deposited on a clean Si wafer. This layer served a dual purpose—it provided a sacrificial release layer and formed the face-down electrode/bond pad place holders. The Al mask was patterned using standard lithography and a wet etchant. The SiO2 was patterned using the Al mask and a dry C2F6 plasma etch (LAM Research TCP9400,Fremont, CA). The Al was removed and the wafer was cleaned after an etch depth of 2.8 μm was achieved.

Figure 1(b). An adhesion promoter, A-174, was dip coated on the wafer surface and rinsed. Next parylene-C was CVD polymerized (Specialty Coating System PDS2010, Indianapolis, IN) on the oxide. Deposition was 4-μm thick.

Figure 1(c). Chemical mechanical polishing of parylene-C was primarily mechanical in nature until the SiO2 was exposed. It is important to minimize shear stress in the parylene film. We discovered the likelihood of film delamination was reduced by using short polish steps (less than 4 minutes), dehydration baking between polishing, minimal carrier down force, and the prior use of an adhesion promoter. CMP was carried out with an IPEC Avanti 472 (Axus Technology, Chandler, AZ) using a KOH silica slurry (0.2 μ, pH = 10, 10% solid load provided by Planar Solutions, Adrian, Michigan). Down pressure and back pressure gradually increased to a maximum of 24.1 kPa (3.5 psi) and 10.3 kPa (1.5 psi), respectively. Parylene thickness was measured using interferometry (Nanospec/AFT, Nanometrics) and the calculated polish rate at full force was 220 nm/min. Between polish runs, wafers were cleaned using a scrubber and NH4OH solution (OnTrak Systems, Fremont, CA) then baked at 90°C for 10 min in a nitrogen purge. Planarization results on a variety of features and spacing were tested (Fig. 2). We achieved 2 μm features and larger with sufficient uniformity. Since our metallization layer was 4200 Å, the 200-1200 Å profile was sufficiently small for this process.

Figure 1(d). An oxygen plasma was used to clean and roughen the parylene-C undercoating (Jun et al. 2001; Pornsin-Sirirak et al. 2002). This plasma cleaning step occurred before each CVD step throughout the process to promote cleanliness and adhesion between layers. The first interfacial parylene layer, PPX-CH2NH2 was polymerized using custom built CVD equipment described previously (Nandivada et al. 2005), although it could be deposited with commercially available equipment as well. The target thickness of PPX-CH2NH2 was 100 nm (80-130 nm typical). The precursor was provided by Joerg Lahann (University of Michigan) with synthesis details described elsewhere (Klee et al. 2004). Peaks at 3361 and 3301 cm-1 in the IR spectrum were used to confirm the presence of the N-H bonds after deposition. Electrode or bond pad openings in the PPX-CH2NH2 layer were created with a photoresist mask and an oxygen plasma etch (Unaxis Plasmatherm 790).

Figure 1(e). Next, metal lift-off (SPR220) was used to pattern the Cr/Au/Cr (100/4000/100 Å). The SPR220 resist surface was treated with tetramethylammonium hydroxide (TMAH) to create an inhibition layer (Redd et al. 1999). Metal was deposited using an e-beam evaporator (Enerjet). The resist was removed with acetone. Following lift-off, wafers were immediately cleaned in fresh acetone, IPA, then spin rinsed for 5 minutes.

Figure 1(f). The second complementary layer of reactive parylene, PPX-CHO, was then deposited (80-130 nm) using the same equipment and parameters as the first layer. Precursor synthesis details are described elsewhere (Lahann et al. 1998; Nandivada et al. 2005). The strong carbonyl stretch at 1688 cm-1 in the IR spectrum was used to confirm the presence of the aldehyde group after deposition. Next, parylene-C was deposited 2.5-μm thick (PDS2010).
Fig. 1

Microfabrication schematic of a parylene-based neural probe. (Cross-section is representative of device in Fig. 3a). (a) High temperature SiO2 deposition. Pattern electrode openings with Al mask and dry etch. (b) Polymerize parylene-C. (c) CMP of parylene-C and SiO2. (d) Polymerize thin layer of reactive parylene, PPX-CH2NH2. Pattern and etch electrode openings. (e) Liftoff metal deposition of Cr/Au/Cr. (f) Complementary layer of reactive parylene, PPX-CHO, and parylene-C. (g) Deposit Ti metal mask for fine structural features. Spin and pattern thick SU-8 structures. Deposit final parylene-C. Spin and pattern thick layer of 9260 resist. (h) O2 plasma etch to define structures. Release devices in BHF. Key features of the design space are highlighted
Fig. 2

Chemical mechanical polishing of parylene over SiO2. Profilometry results immediately after SiO2 is exposed. (a) Profile of an increasing geometric series from 2 to 128 μm of the same track and space followed by the reverse series having wide spaces. Small isolated features achieved the greatest planarization. 2-μm features were achieved. (b) Profile of a 10 and 20-μm feature separated by 30 μm

Figure 1(g). Titanium was deposited 1000-Å thick (evaporator) and patterned (20:1:1 DI:HF:H2O2) to provide a mask during the final oxygen plasma etch of each structure (performed in Fig. 1(h)). SU-8-2025 (Microchem, Newton, MA) was spun on to 40-μm thick and patterned with SU-8 developer (Microchem) to create the core of the probe shank. The SU-8 was cured for 15 min at 150°C. A final parylene-C layer was CVD polymerized (2 μm) which encapsulated the SU-8. We then spun an 80-μm thick AZ-9260 layer (2 step spin) and patterned to create a mask for the thick shank structures. Fine features were achievable on the thin parylene platform (5 μm thickness) using the Ti mask while low resolution features were masked with 9260 resist.

Figure 1(h). The parylene stack was etched using an oxygen plasma to finalize the structural features. 9260 and the Ti mask were removed. Cr on the electrodes and bond pads was also wet etched to expose the Au. Finally, the entire structure was released in buffered hydrofluoric acid. SiO2 removal required about 3 hours then devices were filtered and thoroughly rinsed.

Heat treatment and assembly

Released devices were baked at 140°C for 3 hours to improve adhesion at the surface interfacial boundaries. This heat treatment was previously shown to create an imine bond at the PPX-CH2NH2 to PPX-CHO interface (Chen et al. 2008) and to improve polymer-metal adhesion (Seymour et al. 2009). Neural devices were mounted on a custom PCB employing gold ball studs via thermosonic bonding (Meyer et al. 2001). A protective coating of silicone (MED-4211, Nusil Technology) and epoxy were carefully applied to protect the bond region and PCB.

2.2 Electrode preparation and PEDOT-PSS electrodeposition

Variability in electrode impedance indicated organic residue or metal oxidation was an issue with bare gold electrodes. We used chemical and electrochemical cleaning to lower impedance and reduce variability with a protocol developed for gold electrodes (Rafaela Fernanda Carvalhal et al. 2005). Devices underwent a hot piranha etch for 10 min followed by “electrochemical polishing” in a 0.5 M H2SO4 solution. The polishing protocol consisted of 15 sweeps from -0.1 to 1.2-V vs. SCE (100 mV/s). Finally each device was dipped into ethanol (95%) for 30 minutes before PEDOT deposition.

We selected poly(3,4-ethylenedioxythiophene) (PEDOT) and anion polystyrene sulfonate (PSS) to modify the microelectrodes and lower site impedance. Electropolymerization was carried out in a solution of 0.01 M EDOT monomer (Sigma-Aldrich) and 0.1 M PSSNa (Acros, M.W. 70,000). A current density of 2 pA/μm2 was applied (galvanostatic mode) until a charge density of 1 nC/μm2 was delivered for each site. An Autolab PGSTAT 12 (Eco Chemie, Netherlands) with associated GPES software was used for electropolymerization control.

2.3 Electrochemical impedance spectroscopy

The electrochemical impedance measurements were performed using an Autolab PGSTAT 12 with associated FRA software. The spectral range used for testing interdigitated electrodes (Fig. 5) was 100,000 to 1 Hz and with a single sinusoid at once. All voltage waveforms had an amplitude of 25 mVrms. All impedance measurements were carried out in a Faraday cage in order to minimize external interference. Soak testing was conducted using phosphate buffered saline (PBS, 1X which contains 137 mM NaCl, 2.7 mM KCl, 10 mM Na2HPO4, 2 mM KH2PO4) between 55-65°C for one hour. Heated water was continually monitored and exchanged with water in a large bath surrounding the dish containing PBS, and test device. A two-electrode setup was used for the interdigitated electrodes.

2.4 Edge electrode finite element model

Using COMSOL version 4.0a, we created a 3-D geometry including an insulating substrate, a current sink (modeling a neuron), the extracellular space (modeling tissue), and three different electrodes. Three electrodes were tested: (1) a planar electrode 0.5-μm thick, 15-μm square and positioned 25 μm from the edge, (2) an edge electrode (0.5-μm thick, 7.5-μm wide and 15-μm tall and exposed on both sides, and (3) an edge electrode (5.0-μm thick). We measured the potential at the planar and edge electrodes while moving the spherical current sink to various angles. A neuron was modeled as a 450 μA/cm2 current sink which is the maximum conductance of a Na ion channel, (Lewis 1968)) and positioned 50 μm from the electrode and substrate at increments of 45 degrees (Fig. 6(e,f)). The current sink was 20 μm in diameter (Haug 1987). The tissue was modeled as a large cube (3,000-μm on edge) and was given a conductance of 0.405 S/m (Logothetis et al. 2007). The ground (V = 0) was defined to be the surface of a cube defining the tissue space. The electrode material was assigned a conductance of 8.9e6 S/m (typical for platinum) and defined as having a floating potential. The substrate or shank was 2000-μm long, 100-μm wide, and 5-μm thick. The substrate was defined as an electric insulator. The planar electrode was placed on the surface of the substrate rectangle (Fig. 6(a)) and the edge electrodes were placed inside of a recess at the edge of the substrate (Fig. 6(b-c)) similar to the edge electrodes described here (Fig. 4(c)).

2.5 Acute surgery and recordings

Neural probes were surgically implanted in male Sprague Dawley rats (300–400 g). Anesthesia was administered using intra-peritoneal injections of a mixture of ketamine, xylazine, and acepromazine. The craniotomy spanned approximately 4 mm in the anterior–posterior direction, and 4 mm in the medial–lateral direction and was centered over the M1 and M2 motor cortex (Ludwig et al. 2006). A stereoscope was used to ensure a nearly orthogonal trajectory and avoid any visible surface blood vessels. Immediately before a probe was inserted via a manual stereotaxic drive, the anesthetized subject was moved to a Faraday cage, and then the probe tip was driven to a depth between 1 and 2 mm. All procedures strictly complied with the United States Department of Agriculture guidelines for the care and use of laboratory animals and were approved by the University of Michigan Animal Care and Use Committee.

Recorded neural signals were acquired using a TDT RX5 multi-channel acquisition system (Tucker-Davis Technologies). Neural electrophysiological recordings for all sixteen channels were amplified and bandpass filtered (450-5000 Hz) and sampled at ~25 kHz. Objective classification of high speed neural activity was performed using an algorithm based on an automatic threshold and principal component analysis (Ludwig et al. 2008).

3 Results and discussion

3.1 Parylene-based neurotechnology platform

The CMP-parylene process described allowed us to create neural probes for recording, passive drug delivery, and cell delivery (Fig. 3(a-d)). The SU-8 layer efficiently provides a third dimension to these structures, allowing for either a stiff penetrating tine, or drug reservoir (Fig. 3(a,c,h)). The cell and drug reservoir may be open on one side (Fig. 3(c)) or open on both sides (Fig. 3(d)) by etching the lower parylene layer . An open channel may be used as a reservoir for passive drug delivery by loading with a degradable scaffold. The degradable scaffold would be created separately through a microembossing technique(Yang et al. 2005) and loaded prior to implantation. Built-in electrodes on the bottom of the reservoir may be used to electrically monitor the degradation rate of the scaffold through impedance spectroscopy (Fig. 3(c)). Alternatively, a completely open channel, i.e., no backside, can be used for stem-cell delivery with a dip coating technique (Purcell et al. 2009). These cell-delivery structures now incorporate functional electrode sites. Both of these designs are an alternative to active delivery mechanisms previously employed in neural probes (Papageorgiou et al. 2006; Wise et al. 2008) and may be useful in passive delivery neuropharmaceutical experiments.
Fig. 3

Examples of the design space enabled by this novel parylene-CMP process. (a) Electrodes arranged with 45 micron spacing with front-side, backside, and edge design. The edge electrodes will be tested for improved longevity based on the sub-cellular geometry. (b) Interdigitated electrode used for EIS assessment of interconnect impedance or studying the time course of tissue encapsulation. (c) Open-channel design to be used for passive drug delivery. Channel can be loaded with drug scaffold. (d) Hollow-channel neural probe with 16 channels to be used for cell-based therapies. A wide variety of electrode site options are available: (e) ring electrode, (f) sieve electrode, (g) cantilever electrode, (h) reservoir electrode, (i) large edge electrode, top-view, and (j) PEDOT-coated edge electrode, pseudo-colored SEM, viewed at 45°. 17 × 5 μm footprint. Scale = 200 (a-d) or =25 μm (e-i). Background cropped for contrast (a-i)

The large range in flexibility stems from the ability to create parylene-only structures or add thicker SU-8 structures selectively because it is photodefinable. A parylene stack by itself (5-μm thick) is particularly useful for long interconnects or electrocorticogram arrays (not shown). To estimate the range of flexibility, the equation for the stiffness of a rectangular cantilever is as follows:
$$ k = \frac{{Ew{t^3}}}{{4{l^3}}} $$
where E is the modulus of elasticity for the bulk material, w is the width of the beam, t is thickness, and l the length. Our standard polymer neural probe (Fig. 3(a), 47-μm thick) is estimated to be 0.67 Nm (E = 3 GPa for parylene and SU-8), which is about 15% less than a “Michigan” probe. (14-μm thick, E = 133 GPa for boron doped silicon). Conversely, the same polymer structure without the SU-8 layer is 0.0018 Nm or as a percentage is 0.3%. This indicates that the thin lateral extension does not significantly contribute to the device stiffness despite doubling its width. The 8 mm interconnect structure attached to our standard neural probe is only 55 μNm.

An important goal of this research was to create an edge electrode having a top and bottom surface and the potential for a third surface area (at the edge) through electrodeposition (Fig. 3(i,j)). This feature allowed us to create the smallest footprint (85 μm2) of a functional recording electrode to date. This new electrode design should theoretically increase signal amplitude by reducing the spatial footprint (Moffitt and McIntyre 2005) while achieving a greater surface area than a conventional planar geometry would allow (Fig. 3(j)). In addition to edge electrodes, we provide several examples of single or double-sided electrode geometries (Fig. 3(e-i)) that may also be useful when placed on a thin parylene platform.

3.2 Microelectrode impedance

Previous work has shown that PEDOT can reduce the impedance modulus by nearly two orders of magnitude (Cui and Martin 2003). Our data also showed a significant reduction after PEDOT-PSS electrochemical polymerization. Gold microelectrodes had an average impedance modulus of 6.0 MΩ (at 1 kHz), which was reduced to 210 kΩ after deposition (Table 1). PEDOT electropolymerization reduces impedance by increasing the electrochemical surface area (Fig. 4) (Cogan 2008). The ratio of impedance magnitude drop was greatest on electrodes having the edge geometry (Table 1). This was expected because the exposed edge thickens as more PEDOT is polymerized and effectively creating a third face. The edge thickness increased from 0.4 μm to 1.1 μm after applying a charge density of 1 nC/μm2 measured using electron microscopy (Fig. 4(d)). After applying 2.5 nC/μm2 on another site, the edge was measured to be 1.9 μm, which is an increase of 20% in the geometric surface area relative to the original site size. We expect a thicker edge dimension to be achievable, if desired, since charge densities between 1.3 and 28 nC/μm2 have been reported (Cui and Martin 2003; Ludwig et al. 2006; Abidian and Martin 2008).
Table 1

Electrode Geometry and Impedance (1 kHz)


Au Sitea, MΩ


POST Implantc, kΩ

N = 48

N = 33

N = 15

Planar 17 × 17 μm

5.4 ± 3.7

240 ± 110

380 ± 82

Edge 17 × 5 × 2 μm

6.0 ± 3.1

210 ± 64

460 ± 72

aElectrodes cleaned using piranha etch, H2SO4 polishing, and EtOH

bGalvanostatic deposition, current density = 2.5pA/μ2, charge density = 1nC/μ2

cCleaned in a proteolytic enzyme detergent
Fig. 4

SEM of gold electrodes before and after PEDOT electrochemical deposition. (a) 280 μm2 Au sites. (b) 280 μm2 PEDOT-PSS grown at 2.5 pA/μm2 for a total of 1nC/μm2. (c) 170 μm2 three-sided Au electrode. Thickness was 0.5 μm. (d) PEDOT-PSS on a three-sided edge electrode (same current and charge density). Resulting thickness was 1.1 μm. Images taken at ~45° and 10 keV. Scale = 5 μm

The increase in geometric and electrochemical surface area provided by electrode surface modification is critically important on this size scale. Thermal noise is proportional to the square root of the real part of the impedance, therefore, a surface modification technique is required to decrease the noise level when attempting to discriminate single unit activity. In a few attempts to record single units with bare gold sites of either 280 μm2 or 170 μm2, we found the thermal noise to be more than double the PEDOT modified sites (discussed below). A number of other techniques to increase the electrochemical surface area of an electrode have been reviewed elsewhere (Cogan 2008) and could be used here.

3.3 Interconnect impedance

The electrical stability of the interconnects were also investigated using a high temperature soak test (Fig. 5). Impedance spectroscopy was used to measure the lateral or interfacial impedance of an interdigitated electrode device (Fig. 3(b)) that has no exposed metal and therefore is sensitive only to the capacitive and resistive properties of the dielectric material and interfacial adhesion. The equivalent length of two parallel traces for this design would be 62 mm long (3 mm long, 165 μm wide, 4.0 μm pitch). Unlike the transverse impedance measurement through the parylene film, this measurement is highly sensitive to interfacial delamination (Sheppard et al. 1982; van Westing et al. 1994). Bode plots show the response over one hour of soaking at 60°C and include the dry impedance values as well (Fig. 5). As expected, the dry interfacial impedance was marked by a large modulus and a flat negative phase indicating a purely capacitive response. The impedance modulus declines rapidly at low frequencies once the soak test begins, but still maintains a high value. The decline is asymptotic and the final magnitude at 60 minutes and 60°C remains much higher than the electrode impedance (Table 1). The phase and the high frequency shift of the breakpoint indicate either micro-delamination or pores in the film being filled with water and salts.
Fig 5

(a) One hour soak test results. Bode plots of lateral impedance of an interdigitated electrode at 60°C (N = 3). One-hour soak data indicates these devices initially perform better than parylene-C alone, as expected for a PPX-CHO and PPX-CH2NH2 interface. Does not validate long-term performance. (b) Reprinted from a previous study using a reactive parylene interface (right) compared to a parylene-C only insulation layer (Seymour, Elkasabi et al. 2009). Tests were conducted at 37°C over 60 days also using an interdigitated electrode (N = 5) and show that performance changed little after the first hour

In a previous study, we presented soak data at 37°C for 60 days using the same parylene-C barrier layer and reactive parylene interfacial layers tested here. After one hour the impedance dropped to its approximate minimum value (Seymour et al. 2009) and actually increased slightly over the 60 day period (Fig. 5(b)). Other researchers have estimated that water molecules will diffuse 15-μm/min in parylene (Yasuda et al. 2001). Our results here also show a rapid, asymptotic decline of low frequency impedance as well. While the magnitude of the low frequency change is large, the final modulus remains stable and much higher than the electrode impedance. In the frequency range critical for neural recording, 500-5000 Hz, the impedance modulus decline was small.

3.4 Edge electrode finite element model

The double-sided edge electrode is a unique alternative to conventional planar microelectrodes and has the potential to double the microelectrode density. To achieve a third side or thick, laterally facing edge requires an extra post-fabrication step to electrodeposit a thicker electrode material. One may wonder if a thicker, laterally-facing surface would increase sensitivity to lateral neural activity? Would a thicker electrode increase the viewing angle? A three-dimensional finite element model was created to evaluate the affect electrode position and electrode thickness has on the sensitivity of relative neuron placement. We discuss our finite element model and previously reported microelectrode field models.

A COMSOL model was created to evaluate the electric potential in three-dimensional space (Fig. 6). We measured the signal amplitude on (i) a planar electrode, (ii) a thin (0.5 μm) edge electrode, and (iii) a thick (5.0 μm) edge electrode. An electric field was generated by a current sink (simulating a neuron) that was rotated at a fixed distance from the substrate and incremental angles of 45 degrees (Fig. 6(e,f)). The results (Fig. 6(d)) can be summarized as (i) the thickness of the edge electrode had almost no affect on its recording sensitivity for any neuron position, and (ii) the planar electrode is more sensitive than the edge electrode when the planar electrode is directly facing the neuron but the edge electrode has a wider viewing angle. Since a metal electrode is highly conductive relative to the adjacent tissue it acts as a “float,” which is to say it forces the interface around the electrode to float at the same potential. The “float” effect around the electrode and the relatively small dimensions relative to the spatial gradient of the electric field can be visually seen (Fig. 6(g-i), Supplemental Fig. 1) and illustrates why thickness is not a critical variable to recording sensitivity. The electric field generated by a typical 5-30 μm cell body (Haug 1987) is not greatly influenced by a change in electrode thickness for practical values in the range of submicrons to ten microns. The high sensitivity of a planar electrode has been described elsewhere as an “amplification affect” of the wide substrate which limits the geometric spreading of the electric field (Anderson et al. 2001; Moffitt and McIntyre 2005). Conversely, when a neuron is behind a wide insulating substrate then the signal amplitude is attenuated or “shadowed” which limits its effective viewing angle.
Fig. 6

3D geometries for each electrode type (a) planar, (b) thin edge electrode (0.5 μm), (c) and thick edge electrode (5 μm). A neuron was modeled as a current sink in a conductive medium and held at a fixed distance (50 μm) from the substrate (5 × 100 × 2000-μm). (d) Finite element model summary showing signal amplitude as a function of neuron angle for three electrode types. (e) Top view of planar electrode and seven neuron positions. (f) Edge electrodes (0.5 μm and 5.0 μm thick) and neuron positions. (g,h,i) Electric potential (xy plane) and geometry showing the neuron position with maximal amplitude for each electrode type

The model results suggest that there is no advantage of a three-sided edge electrode, i.e. one with thick electroplating beyond the intention to decrease impedance and therefore reduce thermal noise. The model also suggests a potential tradeoff exists between two or three-sided edge electrodes versus planar electrodes that should be further investigated. If this phenomenon is confirmed experimentally, a wide viewing angle may be advantageous in some applications whereas the larger signal amplitude for fewer neurons may be preferred in others.

3.5 Neural recordings

Acute animal recordings were conducted to evaluate the mechanical robustness and electrical function of the edge electrode design. A polytrode array (45-μm pitch) having electrodes on the face, back, and edge (Fig. 7(a)) was tested for single unit activity and noise levels. Several planar and edge electrodes were tested without PEDOT polymerization and these had noticeably higher noise levels (55 μV peak-peak). PEDOT-modified microelectrodes had an average noise level of 22 μV peak-peak. Single unit activity was isolated with automated principal component analysis (Ludwig et al. 2008). Data was analyzed from 7 insertions in 2 animals using 2 separate arrays. Planar sites had 25 isolatable units on 49 channels while edge electrodes had 57 units on 49 channels during 10 minute sessions. These tests also indicated higher signal amplitudes on the edge electrodes relative to the conventional planar sites (Fig. 7(b, c)). The electrode sites remained functional and with low impedance (Table 1) following multiple insertions.
Fig. 7

Sample in vivo data of planar and edge electrodes. (a) Edge array design has 8 planar sites (4 top-side, 4 back-side) and 8 edge sites on a 45-μm pitch. (b) Snippets extracted using an automated, objective sorting algorithm. (c) High-speed data was filtered from 300 to 5000 Hz. PEDOT was deposited onto all sites (not shown)

4 Conclusion

We present a CMP-parylene process that was successfully employed for the microfabrication of a variety of neural probe designs. A large design space was created most notably by providing a dynamic range of structural stiffness and enabling electrode placement on the front, back, or edge. Preliminary recording data indicates that the small edge electrode design is functional and may be an improved alternative to the conventional planar microelectrode. The implementation of novel edge microelectrode arrays will be used in future testing with the hope of significantly increasing the recording volume and performance over time.


We are grateful to Prof. Joerg Lahann and Dr. Yaseen Elkasabi for providing the reactive parylene dimer and for conducting the CVD polymerization in their laboratory. Ning Gulari provided the critical idea of using CMP in this process and other helpful conversations. Drs. Pilar Herrera-Fierro, Hung-Chin Guthrie, and Ramin Emami shared their considerable CMP expertise which was vitally important. Dr. Kip Ludwig shared his method and software for automatic neural spike sorting. Dr. Mohammad Abidian and Eugene Daneshvar engaged in many helpful discussions regarding conductive polymers. We gratefully acknowledge support from the NIH P41 Center for Neural Communication Technology (EB002030) through the NIBIB.

Supplementary material

10544_2011_9512_MOESM1_ESM.jpg (423 kb)
Supplemental Fig. 1Electric potential slices and geometries from a three-dimensional COMSOL 4.0a model for each combination of neuron position and electrode type. Electric potential (V) shown in xy-plane cutting through the electrode. (a) Planar electrode. (b) Thin edge electrode, 0.5 μm thick. (c) Thick edge electrode, 5.0 μm thick. (JPEG 423 kb)

Copyright information

© Springer Science+Business Media, LLC 2011