of 11
Multiplexed, High Density Electrophysiology with
Nanofabricated Neural Probes
Jiangang Du
1,2,3
, Timothy J. Blanche
4
, Reid R. Harrison
5
, Henry A. Lester
1
, Sotiris C. Masmanidis
1,2,3
*
1
Division of Biology, California Institute of Technology, Pasadena, California, United States of America,
2
Kavli Nanoscience Institute, California Institute of Technology,
Pasadena, California, United States of America,
3
Broad Fellows Program in Brain Circuitry, California Institute of Technology, Pasadena, California, United States of
America,
4
Redwood Center for Theoretical Neuroscience, Helen Wills Neuroscience Institute, University of California, Berkeley, California, United States
of America,
5
Intan
Technologies, Los Angeles, California, United States of America
Abstract
Extracellular electrode arrays can reveal the neuronal network correlates of behavior with single-cell, single-spike, and sub-
millisecond resolution. However, implantable electrodes are inherently invasive, and efforts to scale up the number and
density of recording sites must compromise on device size in order to connect the electrodes. Here, we report on silicon-
based neural probes employing nanofabricated, high-density electrical leads. Furthermore, we address the challenge of
reading out multichannel data with an application-specific integrated circuit (ASIC) performing signal amplification, band-
pass filtering, and multiplexing functions. We demonstrate high spatial resolution extracellular measurements with a fully
integrated, low noise 64-channel system weighing just 330 mg. The on-chip multiplexers make possible recordings with
substantially fewer external wires than the number of input channels. By combining nanofabricated probes with ASICs we
have implemented a system for performing large-scale, high-density electrophysiology in small, freely behaving animals
that is both minimally invasive and highly scalable.
Citation:
Du J, Blanche TJ, Harrison RR, Lester HA, Masmanidis SC (2011) Multiplexed, High Density Electrophysiology with Nanofabricated Neural Probes. PLoS
ONE 6(10): e26204. doi:10.1371/journal.pone.0026204
Editor:
Hiromu Tanimoto, Max-Planck-Institut fu
̈
r Neurobiologie, Germany
Received
August 13, 2011;
Accepted
September 22, 2011;
Published
October 12, 2011
Copyright:
ß
2011 Du et al. This is an open-access article distributed under the terms of the Creative Commons Attribution License, which permits unrestricted
use, distribution, and reproduction in any medium, provided the original author and source are credited.
Funding:
J.D. and S.C.M. gratefully acknowledge support from the Broad Fellowship Program in Brain Circuitry and the Della Martin Fund for Discoveries in
Mental Illness. T.J.B. was supported by NIH-5R21NS066260. H.A.L. acknowledges funding by NIH DA017279. The funders had no role in study design, data
collection and analysis, decision to publish, or preparation of the manuscript.
Competing Interests:
R.R.H. is co-owner of Intan Technologies, which developed the ASIC used in this project. This does not alter the authors’ adherence to all
the PLoS ONE policies on sharing data and materials.
* E-mail: sotiris@caltech.edu
Introduction
Neural probes comprising multiple extracellular microelectrodes
have proven to be an effective tool for recording activity from large
neuronal ensembles [1]. In contrast to most imaging methods,
implantable probes can access virtually any depth of the brain and,
because of their small size, are more conducive to measurements in
awake, freely behaving animals. Much progress has been made in
neural probe technology over the last 6 decades since the pioneering
tungsten microelectrode experiments of Hubel and Wiesel [2,3]. By
scaling up from single to multi-channel electrodes, it is now possible
to record spikes from well over 100 neurons [4], with the number of
simultaneously recorded single-units doubling roughly every 7 years
[5]. However, this apparent wealth of data overlooks an important
limitation of existing technology: densely recording multiple units in
the same region of the brain in a minimally invasive fashion remains
a daunting challenge. This limitation must be overcome in order to
decipher how the brain encodes information across various scales,
from locally connected microcircuits to long-range correlations
between macrocircuits [6]. Instrumentation noise, the rapid spatial
decay in extracellular action potential amplitude, and interference
from more distant co-active neurons, means that electrodes must lie
within
,
100
m
m of the soma to reliably detect and isolate a neuron
[7–9]. There is therefore a strong impetus to develop higher density
electrode arrays to faithfully monitor neuronal subpopulations
within discrete anatomical regions.
Microelectromechanical systems (MEMS) based electrode
arrays are increasingly being used to address this challenge
[10,11]. For example, silicon probes containing dozens of
recording sites on a thin penetrating shaft have yielded important
insights into the function of the hippocampus [12] and visual
cortex [13,14]. However, the development of such devices
inevitably involves a tradeoff between the number of recording
sites, the device width, and the ability to connect the electrodes.
This issue becomes more salient as the number of electrodes per
shaft is scaled up, requiring an increase in the width of the device
to accommodate additional electrical leads (which are also known
as interconnects). Existing neural probes employ
$
1
m
m lead
width and separation [15,16]. We used electron-beam (e-beam)
lithography to reduce these features to sub-micron dimensions.
The resulting high-density lead devices can accommodate a large
number of recording sites without an appreciable increase in probe
width (Table S1).
Performing in vivo very large-scale electrophysiology with
multichannel devices presents an additional challenge: interfacing
probes with the external instrumentation to record neural signals.
Thus, miniaturizing the instrumentation is another crucial
requirement for successfully scaling up neural probe recording
capabilities. The combination of implantable neural inter-
faces with complementary metal-oxide-semiconductor (CMOS)
electronics can fulfill this role [17,18], similarly to what this
advance has done for in vitro microelectrode arrays [19,20]. Using
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a newly developed ASIC, we demonstrate low noise, wide band,
multiplexed measurements with 64 channel nanofabricated neural
probes.
Results
Nanofabricated silicon-based neural probes
Neural probe fabrication was carried out on 100 mm diameter
silicon-on-insulator (SOI) substrates (Figure 1). We developed
three different 64-channel electrode arrays incorporating various
electrode configurations (Figures 2A–C). High-resolution e-beam
lithography was utilized to define nanoscale gold leads and
108
m
m
2
electrode recording pads. The e-beam process produced
leads with a width and spacing as narrow as 290 nm on certain
portions of the array (Figures 2D,E).
Deep reactive ion etching was used to release the probes from
the SOI substrate (Figure 2F) [15,21]. The total shaft thickness was
23
m
m (Figure 2G), and tip widths ranged from 28
m
m for the
linearly configured array in Figure 2A, to 60
m
m for the
‘‘honeycomb’’ array in Figure 2C. These structures were tapered
such that the basal shaft widths were 40 and 85
m
m, respectively.
Detailed array schematics with relevant dimensions are provided
in Figure 3.
Integrated circuitry for signal amplification, filtering, and
multiplexing
An ASIC for performing all essential signal buffering,
amplification, filtering, and multiplexing functions was developed
using 350 nm CMOS process technology (see Table 1 for key
device parameters). The contacts for the 64 input channels were
arranged in a single row on one side of the chip, and their spacing
was matched to those on the neural probe to facilitate wire
bonding on a compact printed circuit board (PCB), displayed in
Figure 4A. The assembled PCB, measuring 9 mm
6
15 mm and
weighing 330 mg after epoxy encapsulation, also contains a
miniature connector, and surface mount passive components for
performing voltage stabilization and division for powering and
digitally controlling the ASIC. The placement of the probe and
accompanying head stage circuitry on the same PCB was
motivated by efforts to minimize the system’s size, weight, and
noise from external interference. Figure 4B illustrates the
operating scheme of a 64 channel neural probe by the ASIC.
Signals were amplified, filtered, and sent through two 32:1 analog
multiplexers before being transmitted off the PCB via a connector
which provided an electrical interface to a data acquisition (DAQ)
device. A minimum of 6 wires (power, ground, channel sweep,
channel synchronization, and the two multiplexer outputs) were
required to perform recording functions with the ASIC; the
remaining connections were required to test electrode impedances
and amplifier operation.
Low noise characteristics of the recording system
To investigate the noise performance of the multiplexed
recording system, we utilized additional circuitry on the ASIC,
which can also perform rapid measurement of electrode
impedance, as well as modification of impedance via electrode-
position (Figure 5A). The noise level markedly decreased with
consecutive gold electroplating cycles (Figure 5B), and asymptot-
ically approached the intrinsic noise of the amplifier (1.7
m
V
rms
,
f = 0.1–6.5 kHz). This trend toward lower noise was strongly
correlated with the accompanying reduction in site impedance
(Figure 5C,D), confirming that gold electroplating is an effective
way to minimize noise of electrodes with small geometric areas
[22]. Figure 5C also indicates that the level of 60 Hz interference
in a relatively unshielded laboratory recording environment
(fluorescent ceiling lights turned on; no Faraday cage) dropped
significantly post-plating, presumably because of the lower
electrode impedance. Although additional interference peaks
appear in the spectrum of gold-plated electrodes, they contribute
,
2% to the total noise power and thus do not significantly disrupt
recording quality. Figure 5E depicts demultiplexed and filtered
segments from extracellular measurements under different imped-
ance and recording conditions, further illustrating the benefits of
electroplating for lowering noise.
Crosstalk and thermal noise in high-density leads
Impedance measurements between adjacent wires yielded
39
6
7M
V
(mean
6
1 s.d. at 1 kHz). Given that the measured
gold-plated electrode impedance values are only
,
1% of this value
(see Materials and Methods), we expect a negligible level of
capacitive cross-talk within these devices. Furthermore, thermal
dissipation in the electrical leads contributes noise to electrophys-
iological recordings. The RMS noise over a bandwidth of
D
f can
be calculated by the expression:
d
V
leads
~
ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi
4
k
B
RT
D
f
p
ð
1
Þ
where k
B
is the Boltzmann constant, R is the resistance of the wire,
and T is the temperature. Resistance measurements across the
length of the probe containing 290 nm leads yielded 9.2
6
0.4 k
V
,
which corresponds to
d
V
wire
= 1.0
m
V
rms
for a bandwidth of
6.4 kHz. This contribution is less than the reported intrinsic noise
from the amplifier and electrode-fluid interface; however, the
fabrication of significantly longer or narrower leads is likely to
produce more noise from thermal dissipation in the wires.
Multiplexed, large-scale data acquisition
Having demonstrated the low noise recording characteristics of
our system, we next evaluated its ability to record neuronal action
potentials. Figure 6 displays a representative epoch of spontaneous
firing activity in the ventral posteromedial thalamic nucleus of an
anesthetised mouse. The majority of functional (63/64) electrodes
in the 1.3 mm long array report spiking activity, much of which
Figure 1. Process flow schematic for the nanofabrication of 64
channel silicon neural probes.
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appears to be correlated between neighboring sites, indicating
that
.
1 electrode often records from the same neuron. The close
electrode spacing is thus well suited for high-density electrophys-
iology with extensive coverage of extracellular fields along the
length of the array. The combination of microvolt level intrinsic
amplifier noise and recording site modification via electroplating
enabled reliable spike detection with peak amplitudes as low as 40
to 50
m
V. Waveforms from a subset of the putative neurons
isolated from the recording in Figure 6 are displayed in Figure 7A.
The maximum spike amplitude per unit was 150
6
80
m
V
(mean
6
1 s.d; Figure 7B), which is in good agreement with other
extracellular recordings with comparable electrode properties
[11,16]. Using a detection threshold of 50
m
V in the 0.4–5 kHz
band, we found that the spike signature of
<
69% of recorded
units could be observed across at least 40
m
m (Figure 7C), which
is equivalent to the average adjacent site spacing on this
particular array. The concurrent measurement of extracellular
action potentials on multiple channels is an effective way to
improve single-unit analysis [9,11]. The inter-electrode distance
on this array appears to offer such signal redundancy for the
majority of measured neurons, but these results suggest the
classification of a small subset of cells could benefit from
,
40
m
m
site spacing.
One of the principal advantages of simultaneously recording
multiple neurons is the ability to assess functional interactions via
cross-correlation analysis [23]. This technique may augment the
analysis of neural circuit information processing [24], and also
provide insights into the synaptic connectivity of local circuits. To
investigate the latter, we measured the spike time cross-correlation
coefficients for 54 thalamic neurons recorded in parallel with the
1.3 mm long array. We compared the correlation coefficient of
each pair of cells to their inter-somatic separation (Figure 7D). The
location of the soma was estimated by the mean spike field position
on the electrode array, weighted by the peak-to-peak spike
amplitude. The figure displays results from 356 of 1,431 total
pairs, which exhibited significant correlations (p
,
0.05) on the
timescale of
6
20 ms. We observed a pronounced increase in the
mean correlation between more proximally located thalamic
neurons, qualitatively consistent with the increased short-range
connection probability observed for a different group of neurons in
cortex [25]. Although neuronal correlations are known to strongly
depend on the behavioral state [24], and may therefore be
disrupted by anesthesia, this analysis demonstrates the potential of
this recording system to link functional and anatomical aspects of
neural circuits with high spatial resolution.
Measurements in freely behaving mice
The low form factor and weight, combined with the reduced
wiring requirements of our multiplexed recording system, makes it
particularly well suited for large-scale electrophysiology in small
freely behaving animals. The 330 mg system described here
weighs
,
1.5% of an adult mouse, representing a substantial
improvement over other systems with comparable numbers of
recording sites, when considering the combined weight of the
electrode and amplifier assemblies [26,27]. We implanted a system
containing a 64 site array into the mouse hippocampus (Figure 8A).
After recovery from surgery, the animal was connected to a
flexible 12 wire tethered cable and allowed to explore a
30 cm
6
50 cm enclosure during acquisition of electrophysiological
data. Current source density (CSD) analysis of LFP signals with a
vertical resolution of
,
28
m
m revealed relatively uniform theta
oscillations in layers between the CA1 and dentate gyrus
(Figure 8B). However, marked shifts in amplitude and phase were
observed on sites around the suprapyramidal and infrapyramidal
blades of the dentate gyrus. We also observed theta phase-
dependent oscillatory firing of hippocampal neurons recorded
Figure 2. 64 channel silicon-based neural probes with nanoscale leads.
(
A
C
) Three configurations of probes. Scale bars, 200
m
m. (
D
) Gold
recording sites and nanoscale leads, patterned with e-beam lithography. The lead width ranges from 1000 nm at the top lefthand corner, to 290 nm
at the bottom righthand corner of the image. Scale bar, 50
m
m. (
E
) A section of interconnecting leads belonging to the ‘‘honeycomb’’ array
configuration displayed in (A) and (D). The narrowest traces have width and spacing of 290 nm. Scale bar, 5
m
m. (
F
) Characteristic chisel-shaped
profile of the probes resulting from deep reactive ion etching of silicon. Scale bar, 25
m
m. (
G
) Silicon device cross-section.
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from the same device during exploratory behavior (Figure 8C).
Neurons were found to be preferentially active in the vicinity of the
negative peak of the theta potential, which is consistent with
previously reported results [28].
Discussion
High-density, large-scale electrophysiology has been an elusive
goal of systems neuroscience. A prerequisite for implementing this
technique is to increase the number of electrodes, and at the same
time miniaturize the instrument to make the technology minimally
invasive and compatible with small freely moving animals. The
electrophysiology system reported here, incorporating a nanofab-
ricated high-density lead probe with an ASIC supporting
multiplexed signal transduction, is one of the smallest and lightest
fully integrated 64 channel recording instruments developed to
date, and the first to include electrode impedance diagnostic and
conditioning functionalities. We have demonstrated our system is
capable of low noise, wide band spike and LFP measurements in
anesthetized and awake behaving mice.
Miniaturization of the interconnecting wires on the probe is an
effective method to increase the number and density of recording
sites without increasing the device’s overall dimensions. Narrower
probes interfere less with surrounding neurons, glia, and the
blood-brain barrier than larger devices. Moreover, it is speculated
that smaller device cross-sections provide favorable conditions for
sustaining stable neural interfaces over longer timescales in chronic
Figure 3. Detailed schematic of the three fabricated probe
configurations.
All probes contain 64 independent recording
channels. Device (
A
) contains 290 nm leads at the most space-limited
area of the array on the base of the shaft. The shaft width is tapered
from 60
m
m at the tip to 85
m
m at the base. The 262
m
m ‘stalk’ indicates
the location at which lead fabrication was transferred from e-beam to a
standard UV lithography process with 3
m
m feature size. Device (
B
)
contains 350 nm leads, 32 sites per shaft, and its shafts are 60
m
m wide.
Device (
C
) contains 500 nm leads, 16 sites per shaft, and its shafts are
tapered from 28 to 40
m
m. All array configurations were fabricated in
two device length variants (2.5 or 5.5 mm long shafts).
doi:10.1371/journal.pone.0026204.g003
Figure 4. A fully integrated 64 channel recording system.
(
A
)
The system, weighing 330 mg, is built on a 9 mm
6
15 mm printed
circuit board containing a nanofabricated neural probe, ASIC, and
miniature connector. The epoxy encapsulating the ASIC is omitted to
show the wire bonds connecting various components on the board.
Scale bar, 1 mm. (
B
) Schematic illustrating the signal transduction path.
Pre-amplification = 100x; amplification = 2x; filter = 1 to 6,400 Hz; multi-
plexing ratio = 32:1. Note that only 6 wires are needed to operate the
system in recording mode; the remaining wires are used for impedance,
amplifier tests, and electroplating.
doi:10.1371/journal.pone.0026204.g004
Table 1.
Measured specifications of the application-specific
integrated circuit used in these experiments.
Parameter
Value
Number of input channels
64
Number of wires needed for recording
6
Size
8.2 mm
6
3.5 mm
Power supply voltage
3.0 V
Power consumption
22.5 mW
Low-frequency cutoff
1.3
6
0.2 Hz
High-frequency cutoff
6,400
6
300 Hz
Amplifier input referred noise (f = 1-6,500 Hz)
2.0
]
V
rms
Amplifier input referred noise (f = 100–6,500 Hz)
1.7
]
V
rms
Gain
194 V/V
Amplifier input impedance (1 kHz)
13 M
V
Average common-mode rejection ratio (60 Hz)
83 dB
Average power supply rejection ratio (60 Hz)
84 dB
Average adjacent input channel crosstalk (1 kHz)
2
84 dB
Note the reduction in number of wires needed for recording relative to the
number of inputs, which is attributed to on-chip multiplexers.
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implants [29,30]. Further work is underway to determine what, if
any, is the maximum cross-section that will avoid inducing
excessive acute and chronic damage. It should be noted that the
mean spike amplitude of units from the hippocampus of awake
mice was less than that of the acute thalamic recordings
(
,
70
m
Vvs150
m
V) and appeared to deteriorate over time.
We speculate this difference might be attributed to increased
neural tissue damage by micromotion of the skull-tethered probe
relative to the brain [30,31]. Thus, in addition to small probe
size, attaining high-amplitude units in chronic experiments may
require a flexible electrical interconnect between the probe and
PCB [32].
The 290 nm leads reported here enabled the fabrication of 64
electrode sites on a single shaft whose width ranged from 60
m
m
near the tip to 85
m
m at the base. This array provided high-density
extracellular coverage across 1.3 mm, which is sufficient to span
all layers of mouse cortex. Extrapolating from these results, leads
would need to be reduced to 58 nm to create a hypothetical 320
site, 85
m
m wide probe that spans the entire length of the mouse
brain along the dorsoventral axis (
,
6.5 mm) with the same density
as the array depicted in Figure 2A. The development of such
narrow wires will contribute additional noise to the measurement,
and presents challenges in fabrication. Capacitive crosstalk is
another important consideration in scaling down of lead
separation, however, given the sub-100 nm thickness of the wires,
parasitic capacitance between adjacent wires is minimal, and can
be counteracted by reducing electrode impedance via electroplat-
ing [33]. Other approaches for maintaining small device profile,
such as integration of multiplexing circuitry directly on the silicon
shaft [18], the use of multiple conductor layers, or alternative
signal transduction paradigms relying on piezoelectric nanoelec-
tromechanical systems [34], are promising but will require some
additional engineering challenges to be resolved before resulting in
viable recording devices.
Figure 5. Low noise performance of the multiplexed recording system.
(
A
) Scanning electron micrographs of 108
m
m
2
electrodes in their
original fabricated state (left), and after gold electroplating to an impedance of 0.15 M
V
. Scale bars, 2
m
m. (
B
) Root-mean-square noise of the
electrophysiological recording system as a function of electroplating time. Measurements were performed in phosphate buffered saline. Tgt. denot
es
plating proceeded until a target impedance of 0.4 M
V
per channel was achieved. Signals correspond to the 0.1–6.5 kHz frequency band. Black points
denote individual channels; red lines represent median value of 64 sites; the blue line corresponds to the input-referred noise floor of the amplifie
rs.
(
C
) Amplifier noise spectral density of a 0 (black), 0.4 (red), 2.2 (green), and 4.9 M
V
(blue) impedance electrical connection to ground. (
D
) Root-mean-
square noise in the 0.1–6.5 kHz band as a function of electrode impedance at 1 kHz. The red line is for illustration purposes. (
E
) Segments of single-
site recordings under various electroplating and environmental conditions. Dashed red lines represent the 99.9
th
percentile range of the voltage.
Unplated and plated electrodes have an impedance of 4.9 and 0.3 M
V
, respectively. The thalamic recording was carried out in a mouse under chloral
hydrate anesthesia. Note that the noise level for a plated probe in thalamus is lower than that of an unplated probe in saline, in spite of background
neural activity.
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Multichannel arrays output dozens of independent signals, and
transferring these data off the recording system for storage and
analysis is not trivial. The conventional approach of connecting
one external wire per channel would result in a prohibitively large
cable bundle tethering the animal to its recording environment.
While ultra-fine wire cables with up to
,
100 wires have been
used, this approach is not ideal for behavioral studies, and
moreover, is unlikely to sustain an additional order of magnitude
improvement in the number of recording channels. On the other
hand, multiplexed electrophysiology with integrated electronic
components is a highly scalable strategy for circumventing the one
electrode per cable requirement. Our ASIC multiplexes input
signals by a factor of 32, and most ASIC functions can be shared
betweeen multiple chips (see Methods). Thus, if a recording system
were to be expanded to 128 channels, the cable needed to record
data would contain 8 wires (plus 8 additional wires for
electroplating/impedance testing). Likewise, readout of an un-
precedented 1,024 channels would require 36 wires, correspond-
ing to well over a tenfold reduction in external wiring
requirements for electrophysiological recordings. We envision
such recordings could be realized via a straightforward extension
of the results presented here, by combining nanofabricated silicon
probes with multiple 64 channel ASIC chips.
In neurophysiological studies of behavior, there is growing
interest in utilizing wireless systems that eliminate the need for
cables altogether [35]. However, power requirements invariably
force a tradeoff between weight, measurement frequency band-
width, digital sampling resolution, electronic noise, and number of
channels in the wireless transmitter [36]. Current state-of-the-art
wide-band, multichannel wireless neural interfaces approach or
exceed the size and weight of an adult mouse [37,38]. Our
lightweight wired system affords low noise measurements of spikes
and local field potentials, a feature which is absent from some
wireless instruments. It is likely that significant advances occur in
wireless technology are on the horizon; but until then, wired and
multiplexed recording technology will remain particularly attrac-
tive for behavioral studies of mice, which are amenable to a broad
spectrum of genetically encoded brain circuit dissection techniques
[39].
Finally, the wealth of data emerging from high-density
multichannel probes presents new challenges and opportunities
for data analysis [5,40]. Significant headway has already been
made automating spike classification and sorting routines [41,42].
Additional work is needed to automatically and reliably extract
single-unit spike times from electrode array measurements
comprising an arbitrary number of contiguous input channels,
and receiving extracellular signals from several co-active neurons.
Materials and Methods
Probe nanofabrication
Silicon probe fabrication was carried out on a 100 mm diameter
SOI substrate, with a buried oxide (BOX) layer of 1.2
m
m, and
device layer of 20
m
m. The wafer was initially placed in a thermal
oxidation furnace at 1000
u
C to produce a silicon oxide film which
was matched in thickness – and hence in stress to the BOX layer.
Due to the relatively long time involved in e-beam lithography, we
patterned small (
#
3
m
m) features and the 108
m
m
2
electrode pads
Figure 6. Parallel recording capabilities of the multiplexed electrophysiology system.
Traces are filtered from 0.3–5 kHz to highlight
spiking activity, and are plotted near their corresponding location on the shaft of the probe displayed on the left (for each horizontal left-right pa
ir of
pads, the trace for the left-hand pad appears just above the trace for the right-hand probe). Measurements were made in the mouse thalamus with a
probe with post-plating impedances of 0.6
6
0.4 M
V
(mean
6
1 s.d.). The majority of sites report action potentials, and only one of the 64 sites
(uppermost trace) is non-functional.
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with e-beam lithography, and larger features with conventional
UV contact lithography. Metal was deposited after each of these
steps. Note that the width of the interconnecting leads was only
290 nm on the areas of the array where space was most limited
(for example, see Figure 2E). In other areas with more space for
routing wires, the width was increased to minimize the likelihood
of fabrication-related device defects.
For the e-beam process, a polymethyl-methacrylate (PMMA)
resist layer was spin-coated and baked before exposing the
nanoscale leads in an EBPG 5000
+
writing tool (Leica). Patterning
across the wafer containing 170 devices required approximately
7 hours of e-beam writing time on this instrument. After
developing the exposed resist, the wafer was transferred to a
thermal evaporator for deposition of the first metallic layer (3 nm
Cr for adhesion and 70 nm Au). The unpatterned resist and metal
were lifted off by immersing the wafer in acetone, briefly applying
ultrasonic agitation, and rinsing in isopropyl alcohol. For the
second metallization layer involving larger electrically conducting
features located at least 400
m
m from the narrow front portion of
the shaft, a standard contact UV photolithography process was
employed. This involved spin-coating a negative tone photoresist
(AZ nLOF2020) and baking it onto the wafer, after which the
wafer was transferred to mask aligner (MA6, Karl Suss). Patterns
belonging to the second metal layer were aligned, with an
accuracy of
,
0.8
m
m, to alignment marks created during the e-
beam lithography step. After resist development, a 3nm Cr/
100nm Au layer was thermally evaporated, followed by lift-off in
acetone.
Next a 2.2
m
m parylene-C layer was coated onto the front side
of the wafer; electrode recording sites and wire bond contacts were
then exposed by etching the parylene film in O
2
plasma, using
photo-patterned resist as the masking material. The thermal oxide
material on the front of the wafer was then similarly masked and
etched. Subsequently, the silicon device layer was anisotropically
etched via deep reactive ion etching (DRIE, Bosch process), again
using the photoresist as masking material. The exposed BOX layer
was then etched down to
,
100 nm but not completely removed.
The wafer was flipped over and mounted onto a temporary carrier
wafer with wax or photoresist. The bottom oxide was completely
removed, and XeF
2
, which has high etching selectivity for Si over
oxide, was applied to remove the 350–400
m
m SOI Si substrate
layer. The exposed 100 nm of buried oxide attaching the probes to
the remaining SOI frame was then carefully etched with CF
4
plasma. Finally, the finished probes were released from the
Figure 7. Nanofabricated probes facilitate high-density elec-
trophysiological recordings.
(
A
) Representative waveforms corre-
sponding to putative single neurons. Gray lines denote the mean
waveform. (
B
) Distribution of the maximum peak-to-peak amplitude
per unit. The cutoff for reliable spike detection was 50
m
V. (
C
)
Cumulative distribution of the maximum distance across which a unit
could be detected (defined as a mean spike-triggered waveform
.
50
m
V). 31% of units were observed on only a single site; the
remainder were measured on
$
2 recording sites, which have a nearest
neighbor spacing of
,
40
m
m. (
D
)Spiketimecross-correlation
coefficient from 54 simultaneously recorded putative thalamic units,
as a function of estimated inter-somatic separation.
doi:10.1371/journal.pone.0026204.g007
Figure 8. Electrophysiology with nanofabricated probes in
awake behaving mice.
(
A
) Nissl-stained brain section overlaid with a
schematic of the probe at its stereotaxically implanted location. Each
silicon shaft is 60
m
m wide. Scale bar, 200
m
m. (
B
) Current source
density analysis of local field potentials across the hippocampus with a
vertical resolution of
,
28
m
m. Measurements correspond to the right-
most shaft, which is co-localized with both the supra- and infra-
pyramidal blades of the dentate gyrus (DG). The data were gathered
during home cage exploratory behavior. The CSD is normalized to
6
1.
Scale bar, 100 ms. (
C
) Waveforms of two putative single units recorded
from this probe across all sites on their entire respective shafts, with
histograms showing theta phase locking of spikes. Dashed ellipses
indicate the sites exhibiting the highest extracellular action potentials
for these units. Measured theta oscillations on top of the histograms are
for reference. Theta oscillations were measured from the upper right-
most electrode near the CA1 pyramidal cell layer.
doi:10.1371/journal.pone.0026204.g008
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October 2011 | Volume 6 | Issue 10 | e26204
temporary carrier with acetone, and rinsed with isopropyl alcohol,
ethyl alcohol and deionized water. Devices were created with a
shaft length of either 2.5 or 5.5 mm. We found that
,
40% of
probes survived the entire process with
#
5 defective channels. The
majority (
.
70%) of defects occurred during the e-beam
lithography and subsequent metal liftoff step. We foresee
opportunities for improving device yield through further research
and refinement of the e-beam lithography process.
Recording system assembly
Custom-designed PCBs were initially assembled with surface
mount resistors and bypass capacitors that perform voltage
stabilization and division (5 V to 3 V) for powering and digitally
controlling the ASIC. A connector (nano-series, Omnetics) was
also soldered onto the PCB. The ASIC was diced to a rectangular
shape of 8.2 mm
6
3.5 mm, and thinned to 250
m
m prior to
assembly. The neural probe and ASIC were glued onto the PCB
with a drop of PMMA dissolved in anisole. A total of 64 wire
bonds were then created to connect the probe to the ASIC inputs;
an additional 25 bonds were made from the ASIC to the PCB.
The electrical contact areas were then encapsulated in opaque
epoxy to prevent damage and photoelectric artifacts. Drawings are
available from the authors for the PCB and silicon probe designs.
For electrode impedance measurements and gold electrodepos-
tion, the system was connected to a DAQ device (USB-6251,
National Instruments) via an intermediate PCB containing
circuitry for performing impedance checking and conditioning.
The circuitry on this intermediate PCB consisted of mechanical
relays for rapidly connecting or disconnecting the reference
electrode from ground in order to respectively allow impedance
checking or electroplating. In addition, a bank of surface mount R-
C components allowed generation of a 1 nA sinusoid from a
1 kHz 5 V square wave, and was used as the bias signal during
impedance measurements. Finally, resistors divided a 5 V DC
signal to
2
1.5 V between inputs and reference for electroplating.
The ASIC can accommodate a range of other settings, for
example, DC potentials of up to
6
3 V for electroplating. All
inputs to the intermediate PCB were supplied by the same DAQ
device used in recording.
After attaining the targeted impedance values, the ground and
reference pins of the ASIC were permanently shorted, and a
stainless steel ground wire was soldered onto the PCB. For
neuronal recordings the system shown in Figure 4 was connected
directly to the USB-compatible DAQ device via a flexible 1 meter
cable (consisting of 12, 36 gauge wires) plugged into the connector
on the PCB. The integrated multiplexers permitted sampling
speeds of up to 31.25 kilosamples per second (kSps) per channel.
The experiments reported here were carried out at 22 kSps per
channel, corresponding to the upper aggregate sampling rate of
the DAQ device, which multiplexed signals with 16-bit resolution
before sending it to a computer for storage and offline analysis.
The cable provided the following inputs/outputs to the system:
(1) 3 V battery-supplied power source; (2) ground, (3 & 4) analog
output of the two multiplexers; (5) digital input channel selector,
commanding the ASIC to sample the next channel when this line is
activated; (6) digital channel indicator (synchronization), which
reports a high voltage when the first input channel in the
multiplexing sequence is active (i.e., Ch. 1 and 33 for the 1
st
and
2
nd
multiplexer, respectively); (7) digital channel reset, which
commands the ASIC to return to the first input channel; (8) digital
settle, for rapidly discharging all capacitors in the front end
amplifiers to ground in the event of amplifier saturation; (9) digital
test enable switch, for toggling between recording and impedance
test (i.e., check/conditioning) mode; (10 & 11) input voltages for
performing impedance tests; and (12) bank selection switch, for
toggling between the 1
st
and 2
nd
multiplexer during impedance
tests. During recording only pins 1 through 6 were used; pins 7
through 12 were either shorted to ground or disconnected from the
DAQ device. All 12 pins were connected during impedance
measurements, although pin 8 was never activated. The system was
operated with software created in LabVIEW (National Instru-
ments), and data was analyzed offline with Matlab (MathWorks). To
convert amplified signals to input-referred voltage we divided them
by 200, which corresponds to the nominal gain of the amplifiers
(actual measured gain = 194). Note that most ASIC functions (pins
1, 2, 5–9, 12) can be shared across multiple chips to maintain low
wiring requirements for even greater numbers of channels.
Electrode plating
The probe was immersed in gold plating solution (Sifco), and a
2
1.5 V DC potential was applied through the 11
th
control
terminal on the active channel relative to a platinum wire
reference. We used mechanical relays on the intermediate PCB
described above to rapidly alternate between electroplate and
impedance check modalities with the ASIC. Channels were
selected by controlling the 5
th
and 12
th
ASIC terminals described
above. For electrophysiological recordings, electrodes were plated
until their impedance fell into the range of 0.4 to 0.8 M
V
unless
stated otherwise in the main text. If the impedance would not drop
below 3 M
V
, that recording site was classified as faulty and
excluded from further analysis. To minimize the likelihood of
short-circuiting adjacent recording sites by excessive electroplat-
ing, the impedance was not lowered below 80 k
V
.
We observed a small increase in the impedance of electroplated
channels after an in vivo experiment, which was likely due to
detachment of a portion of the electrodeposited gold material
during insertion. For example, after one acute experiment the
probe’s impedances rose from 0.5
6
0.3 M
V
to 0.8
6
0.4 M
V
,
corresponding to an increase of 0.3
6
0.3 M
V
per channel (values
denote mean
6
s.d.).
Electrode impedance measurements
Impedance was measured by feeding a 1 nA, 1 kHz sinusoidal
current into the ASIC, and reading out resulting the voltage signal
from the multiplexer. Impedance was calculated as the ratio of AC
voltage to current. Values were confirmed by comparing measure-
ments to a commercial impedance measurement device (nanoZ,
White Matter LLC). To obtain good correspondence with the
nanoZ, a correction factor of 0.62 had to be applied to the V/I
ratio as measured by the ASIC and accompanying instrumenta-
tion. All impedance measurements reported here were made in
phosphate buffered saline (PBS).
Noise measurements
The probe was immersed in PBS solution, and 200 s duration
multiplexed measurements were carried out on 64 channels with
,
22 kHz sampling rate per channel (exact value: 1/44.8
m
s for all
experiments). Root-mean square noise values were obtained by
taking the standard deviation of each channel’s voltage-time trace
after filtering from 0.1–6.5 kHz with a 3
rd
order Butterworth filter.
No additional signal conditioning, e.g., notch filter was applied. The
noise spectral density (NSD) was obtained from the raw demulti-
plexed data using the Welch periodogram estimation method.
Amplifier gain and phase measurements
The nominal gain and phase response of the amplifier was
measured by applying sinusoidal voltage signals from a benchtop
Nanofabricated Neural Probes
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8
October 2011 | Volume 6 | Issue 10 | e26204