of 6
Nanofabricated Neural Probes for Dense 3
D Recordings of Brain Activity
Gustavo Rios,
Evgueniy V. Lubenov,
Derrick Chi,
§
Michael L. Roukes,
*
,
,
,
§
,
and
Athanassios G. Siapas
*
,
,
Division of Biology and Biological Engineering,
Division of Engineering and Applied Science,
§
Division of Physics, Mathematics,
and Astronomy, and
Kavli Nanoscience Institute, California Institute of Technology, Pasadena, California 91125, United States
*
S
Supporting Information
ABSTRACT:
Computations in brain circuits involve the coordinated
activation of large populations of neurons distributed across brain areas.
However, monitoring neuronal activity in the brain of intact animals
with high temporal and spatial resolution has remained a technological
challenge. Here we address this challenge by developing dense, three-
dimensional (3-D) electrode arrays for electrophysiology. The 3-D
arrays constitute the front-end of a modular and con
fi
gurable system
architecture that enables monitoring neuronal activity with unprece-
dented scale and resolution.
KEYWORDS:
Nanofabricated neural probes, high-density microelectrodes, 3-D scalable packaging, brain mapping
B
rain functions such as perception, motor control, learning,
and memory arise from the coordinated activation of neu-
ronal assemblies distributed across multiple brain areas. While
major progress has been made in understanding the response
properties of individual cells, circuit interactions remain poorly
understood. One of the fundamental obstacles to under-
standing these interactions has been the di
ffi
culty of measuring
the activity of large distributed populations of neurons in
behaving animals.
1
5
Electrophysiology has been the gold stan-
dard for monitoring the brain because it measures the electrical
activity of neurons directly and at a high temporal resolution,
su
ffi
cient to capture in detail even the fastest neuronal events.
The main drawback of electrophysiology has been the inva-
siveness of the recording electrodes and the consequent limits
on the spatial extent and spatial resolution of the obtained
signals.
Research on electrical probes has focused on overcoming
these challenges by scaling up the number of recording sites
while minimizing their invasiveness.
1
,
4
,
6
15
These are inherently
competing objectives because smaller probes, with mechanical
dimensions that minimize tissue displacement, o
ff
er less surface
area and volume for electrode sites, interconnects, and active
circuit elements.
16
Furthermore, as electrode count increases,
so does the need to bring active signal conditioning and multi-
plexing components closer to the brain, as the number of
passive interconnects exceeds the limits of connector and teth-
er cable density. This, in turn, introduces another dimension
to the invasiveness of the recording system
the amount
of electrical power it dissipates as heat into the brain tissue.
Chronic viability of the probes imposes additional constraints
Received:
June 28, 2016
Revised:
September 17, 2016
Published:
October 21, 2016
Figure 1.
Recording system modules. (a) Examples of six realized
neural probe designs. The number of shanks (1
8), intershank
spacing (250
1000
μ
m), recording site arrangement, and pitch (20
65
μ
m) are con
fi
gurable. All designs support 256 electrodes per layer,
connected to a standard 16
×
16 interconnect matrix with 200
μ
m
pitch at the probe base. (b) Designs of two ultra
fl
exible cables
(fabricated on either 10
μ
m thick Parylene C or 15
μ
m thick
polyimide) used to interface the neural probes to the signal
conditioning PCB. (c) Designs of two di
ff
erent signal conditioning
PCBs (headstages). Each performs
analog signal conditioning,
multiplexing, and digitization of 256 analog inputs. The top circuit
(acute) measures 39
×
37 mm, weighs 4.5 g, employs 8 Intan
RHD2132 QFN packaged chips, and requires 8 output LVDS lines (16
wires), while the bottom circuit (chronic) measures 30
×
32 mm,
weighs 1.2 g, uses 4 Intan RHD2164 bare dies, and requires four
output LVDS lines (eight wires). (a
c) Scale bar: 3.4 mm.
Letter
pubs.acs.org/NanoLett
© 2016 American Chemical Society
6857
DOI:
10.1021/acs.nanolett.6b02673
NanoLett.
2016, 16, 6857
6862
This is an open access article published under an ACS AuthorChoice License, which permits
copying and redistribution of the article or any adaptations for non-commercial purposes.
on the biocompatibility of all materials that come in direct
contact with brain tissue as well as on the
fl
exibility of the probe
itself and its coupling to the rest of the system.
17
21
Finally,
relating the measured extracellular potentials to the underlying
circuit elements requires solving an inverse problem to obtain
a detailed current source density (CSD) distribution.
22
The
quality of this CSD estimate critically depends on the density of
electrodes and their three-dimensional (3-D) arrangement on a
regular array of known dimensions and relative position to the
tissue.
4
While signi
fi
cant progress has been made in solving the
above issues individually, addressing them simultaneously
within a full system has remained a challenge.
Here we describe the development of a modular, scalable
system for dense 3-D chronic electrophysiology that addresses
many of the challenges above. The front end of the system
is comprised of passive high-density nanofabricated neural
probes (
nanoprobes
,
Figure 1
a)
2-D arrays of minimally inva-
sive shanks with nanoscale interconnects
that are subse-
quently stacked into a 3-D array of precise geometry with over
a thousand recording sites. The front end of the system is
mechanically and thermally
decoupled from all active
components through high-density
fl
exible cables (
Figure 1
b),
which interface the neural probes to the signal conditioning,
multiplexing, and digitizing circuitry. The latter is housed on
compact, lightweight PCBs (Sierra Circuits, HDI PCB technol-
ogy), compatible with acute and chronic experimentation
(
Figure 1
c). We describe the design, fabrication, and assembly
of the system and its performance characteristics. We also
demonstrate the realized yield and quality of electrophysio-
logical recordings in experiments with awake head-
fi
xed mice.
Each neural probe is a thin (21
μ
m) silicon device with a
square base (3.4
×
3.4 mm) and up to eight narrow (65
μ
m)
shanks containing a total of 256 microelectrode sites (8
×
16
μ
m ovals) distributed in single or double row con
fi
gurations
(
Figures 1
2
). The base houses a 16
×
16 interface matrix of
100
μ
m circular pads with 100
μ
m edge-to-edge spacing, which
constitutes the standardized interface between the probe and
the rest of the system (
Figure 1
a). In order to minimize the
invasiveness of the shanks, while maintaining high electrode site
density, the following design choices were implemented. First,
the width of shanks was kept at a minimum in order to reduce
mechanical invasiveness through tissue displacement.
23
In all
but one design, the maximal shank width in the span containing
electrodes was less than 65
μ
m(50
μ
m on average), while shanks
Figure 2.
Minimally invasive high-density neural probes. (a) Microscope images of four di
ff
erent shank tips with di
ff
erent electrode con
fi
gurations.
The shank width at the electrode furthest from the tip is less than 65
μ
m for all but one design, shown in g, while shank width at the base is 100
μ
m.
(b) Shank width is minimized by using nanoscale interconnects. Shank areas subjected to sectioning by focused ion beam (FIB) milling are marked
with black lines, and the red rectangle marks a region imaged with scanning electron microscopy (SEM). (c
f) False color SEMs indicating di
ff
erent
materials according to color legend on the right. (c) Shank cross section reveals nanoscale (300
×
300 nm) copper interconnects (orange) with a
pitch of 600 nm buried in 1.6
μ
m of oxide insulation (purple, see
Supplementary Figure S1
for details). (d) Cross section at the shank edge
demonstrates conformal coverage of shank sidewall by a biocompatible Parylene HT layer (tan). (e) Two gold electroplated microelectrodes
(yellow) demonstrate the increase in electrode surface area and roughness while preserving planar dimensions. (f) Tip of the 21
μ
m thick shank
demonstrates conformal coverage of three sides (top and sidewalls) by Parylene HT (tan). The bottom side of the shank is composed of 900 nm
SiO
2
, which is also biologically inert. (g) SEM and stereoscope (inset) image of a probe mounted on a slightly wider silicon spacer. The probe
thickness is 21
μ
m throughout and can be assembled onto a spacer of arbitrary thickness to control the pitch of a 3D stack (300
μ
m thick spacer
shown). (h) Devices fabricated at a commercial foundry (LETI, Grenoble, France) on 200 mm SOI wafers (inset) are fully released and anchored in
place on an SOI wafer.
Nano Letters
Letter
DOI:
10.1021/acs.nanolett.6b02673
NanoLett.
2016, 16, 6857
6862
6858
where much narrower near the tip (24
μ
m) and only gradually
widened to about 100
μ
m near the probe base (
Figure 2
a).
Narrow shanks were made possible by utilizing nanoscale
interconnects, which had a 300
×
300 nm cross-section and
were spaced at 300 nm (
Figure 2
c). Second, electrodes were
small in area (117
μ
m
2
) and shaped as ovals elongated parallel
to the shank axis (8
×
16
μ
m), which further minimized the
shank width (
Figure 2
b). Low impedance was achieved in this
small microelectrode area by gold electrodeposition), which
increases the e
ff
ective electrode surface area without altering its
planar dimensions (
Figure 2
e).
24
,
25
Third, shanks were coated
with a Parylene HT biocompatibility layer
26
on 3 sides, while
the backside was made of biologically inert glass (silicon oxide)
(
Figure 2
d,f). Fourth, the probes are completely passive devices
interfaced to all powered electronics through a 15
μ
m thin
ultra
fl
exible cable (Metrigraphics LLC), which isolates the
probes both thermally and mechanically from the rest of the
system. Finally, while all devices were developed in-house on
100 mm SOI wafers using electron beam lithography and
MEMS fabrication procedures (Kavli Nanoscience Institute,
Caltech), the
fi
nal probes were nanofabricated using
a hybrid CMOS/MEMS process on 200 mm SOI wafers at a
commercial state-of-the-art semiconductor foundry (LETI, Gre-
noble, France; see
Supplementary Figure S1
for details). This
improved device yield, quality, and consistency (
Figure 2
h).
While recent work has highlight
ed the potential advantages
that more
fl
exible substrates may provide,
27
31
we fabricated
the neural probes using silicon on isolator (SOI) wafers with thin
(17
μ
m) device layer in order to guarantee precise and
reproducible three-dimensional (3-D) electrode arrangements
(
Figure 3
). Mechanical decoupling of the probe was achieved by
interfacing it to the rest of the system using ultra
fl
exible cables.
The probe, cable, and PCB were
fl
ip-chip bonded together
(Fineplacer Lambda, Finetech) u
sing the anisotropic conductive
fi
lm (ACF, H&S Hightech, TCF1051GY for probe to
fl
ex cable
bond; TGP2050N for
fl
ex cable to PCB bond) to produce a fully
functional 2-D recording module (
Figure 3
a). The use of ACF
was essential for accomplishing
low contact resistance (<1
Ω
)
connections within the compact,
fi
ne-pitched probe pad matrix.
The2-Dmoduleswereusedaslayersthatwerethencombined
together into the 3-D stack (
Figure 3
b). The neural probes
comprising the 3-D electrode arra
y were precisely aligned with the
fl
ip-chip bonder, spaced using silicon spacers of 300
μ
mthickness
(
Figure 2
g), and bonded together with either polyethylene glycol
(PEG, MW: 3000, Sigma; temporary bond) or thin epoxy sheets
(AiT Technology, ESP8680-HF; permanent bond;
Figure 3
c).
Notice that the 3-D electrode array is highly con
fi
gurable through
choice of neural probe model, spacer thickness, and probe
alignment. To demonstrate the power of this approach we
assembled a dense 3-D electrode array with 1024 electrodes
spanning a 0.6 mm
3
volume (
Figure 3
d).
Figure 3.
Recording modules con
fi
gured as a 3-D array with 1024 electrodes. (a) Acute (left) and chronic (right) 256-channel recording modules
consisting of a neural probe,
fl
exible cable, and signal conditioning PCB. (b) Four recording modules are assembled as layers into a stack to form a
1024-electrode 3-D array (system weight 20 g, including 3 mm tall PCB brass spacers; chronic system weight 6.8 g). (right) Close-up view of the
stacked neural probes. (c) 3-D electrode array is highly compact and con
fi
gurable. The shank spacing of the selected neural probe controls electrode
pitch along the
x
-axis, with available options ranging from 250
μ
m to 1 mm. Electrode spacing along the shanks of the selected neural probe controls
pitch along the
y
-axis, with available options ranging from 12 to 65
μ
m. The silicon spacer thickness (arbitrary) controls the electrode pitch along the
z
-axis. A minimum
z
-pitch of 50
μ
m, which can be achieved without the use of the spacer, is determined by the combined thickness of the neural
probe base (21
μ
m), ACF (14
μ
m), and
fl
exible cable (15
μ
m). (d) The 3D electrode array used to obtain in vivo recordings. Its
x
y
z
pitch is
250
12
350
μ
m, and the volume enclosed by the array is 750
756
1050
μ
m, giving an electrode density of 1024 electrodes for 0.6 mm
3
.
Nano Letters
Letter
DOI: 10.1021/acs.nanolett.6b02673
NanoLett.
2016, 16, 6857
6862
6859
Our system architecture separates the active signal
conditioning circuits from the neural probe to minimize heat
dissipation from the active electronics into the brain (see
Supplementary Figure S3
for details). This requires careful
budgeting of parasitic capacitances and electrode impedances
in the overall system design (
Figure 4
). Cross-talk between
adjacent traces grows with electrode impedance (
Figure 4
b), so
we used gold electrodeposition to increase the e
ff
ective micro-
electrode surface area thereby lowering impedance by an order
of magnitude (
Figure 4
a). The electrochemical impedance
spectra obtained before and after plating allowed us to esti-
mate the parameters of the equivalent circuit representing the
electrode
electrolyte interface (
Figure 4
a) and to map out the
crosstalk dependence on frequency and electrode character-
istics (
Figure 4
b). This analysis demonstrates that electrode
impedance below 0.5 M
Ω
(0.3 M
Ω
) at 1 kHz limits cross-talk
to values below 1% for all frequencies below 1 kHz (10 kHz),
respectively. This range of microelectrode impedance values
could be readily achieved by gold electrodeposition. Our anal-
ysis of the impact of coupling capacitance between adjacent
traces on cross-talk (
Figure 4
c) in
fl
uenced our design choice of
the nanoscale interconnect cross-section, spacing, and total
length. With these considerations, we achieved low cross-talk
and system noise of 4.8
μ
V (9.4
μ
V) RMS measured in saline
for plated (unplated) electrodes, respectively (
Figure 4
d), in a
3-D electrode array with unprecedented density (
Figure 4
e).
In order to experimentally validate the system, we recorded
electrophysiological activity from the hippocampus of awake,
head-
fi
xed mice, using the 1024 electrode 3-D array (
Figure 5
,
see
Supplementary Figure S2
for data acquisition details and
Figure S4
for experiment setup details). The raw broadband
extracellular signal from one layer of the 3-D stack is shown in
Figure 5
a. Notice that the low frequencies of the broadband
signal, known as the local
fi
eld potential (LFP), display clear
and systematic spatiotemporal variations, which are a prominent
and recognizable feature of hippocampal activity (
Supplementary
Figures S6 and S7
). In contrast, the high frequency band con-
tains high amplitude spikes, spatially restricted to nearby micro-
electrodes that were anatomically close to the pyramidal cell layer
(
Figure 5
a,b). Furthermore, the same spikes are clearly seen on
multiple neighboring recording sites, thereby allowing for
successful triangulation of the source neuron and spike sorting
(
Figure 5
b,
Supplementary Figure S8
). Although, the relative
location of the 3-D electrode array with respect to the hippo-
campal circuitry can be inferred from the recorded patterns of
electrophysiological activity alone, we directly veri
fi
ed it through
histological sectioning and analysis (
Figure 5
c,
Supplementary
Figure S5
).
One key objective of brain activity mapping is the ability to
observe the
fi
ring of all neurons within a brain volume. How
close does the dense 3-D electrode array described here bring
us to achieving this ultimate goal? Because action potential
amplitudes decay rapidly in the extracellular space, each site
can only detect spikes originating within a sphere of radius
R
100
150
μ
m, centered at the electrode. The union of
these spheres, one for each electrode site, gives the observable
Figure 4.
System characteristics and comparison to other 3-D neural recording systems. (a) Equivalent circuit model (top inset) for an unplated
(red) and gold-plated (blue) electrode
electrolyte interface derived from electrochemical impedance spectroscopy (EIS) data, displayed as Bode
plot and captured using a fully passive assembly. Equivalent circuit parameters for unplated (plated) electrodes were: spreading resistance
R
s
=20k
Ω
(15 k
Ω
), charge transfer resistance
R
ct
=55G
Ω
(89 G
Ω
), constant phase element (CPE) exponent
α
= 0.88 (0.91), CPE prefactor
Q
=60
×
10
12
(750
×
10
12
)s
α
/
Ω
, resulting e
ff
ective capacitance
C
e
= 9.4 pF (243 pF).
C
cell
is the parasitic capacitance introduced by the measurement setup,
C
cell
= 12 pF. Notice that gold electroplating reduces the electrode impedance by an order of magnitude (from 3.8 M
Ω
to 500 k
Ω
at 1 kHz) due to a
corresponding increase in the electrode
se
ff
ective double layer capacitance. Microscope images (bottom inset) of an unplated (red) and plated
(blue) electrode. Scale bar: 8
μ
m. (b) Equivalent circuit model (inset) used to analyze crosstalk between two adjacent interconnects.
10
Traces
correspond to increasing electrode impedance (values at 1 kHz shown on right) while all remaining parameters are kept constant at values estimated
for our system (coupling capacitance between adjacent traces,
C
ss
= 1.35 pF, trace shunting capacitance to ground,
C
sh
= 2.5 pF, ampli
fi
er input
capacitance
C
L
= 12 pF). (c) Cross-talk at 1 kHz for increasing values of the coupling capacitance (traces as in b). Notice that, even for low
impedance electrodes, a coupling capacitance above 8 pF results in crosstalk in excess of 1%. (d) System noise (RMS) of unplated (left, 9.4
μ
V
median) and plated (right, 4.8
μ
V median) microelectrodes (bandwidth: 0.1 Hz to 7.5 kHz). The input referred noise of the ampli
fi
er is 2.4
μ
V.
(e) Electrode count (1024) and density (1720 el/mm
3
) of our realized 3-D array in comparison with previous work.
Nano Letters
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DOI:
10.1021/acs.nanolett.6b02673
NanoLett.
2016, 16, 6857
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volume. The fraction of observable neurons can then be
estimated as the ratio of the observable volume to the total
array volume. For the dense 3-D array, this ratio initially grows
approximately as the second power of the sphere radius and the
estimated fraction of observable neurons is 42% (60%) for
R
=
100
μ
m (125
μ
m), respectively. In contrast, the volume of
tissue displaced by the array is less than 1%.
The ability to detect spikes is only a necessary condition for
successfully isolating the
fi
ring of a source neuron. In addition,
spikes from the source neuron should be detected with suf-
fi
cient amplitude on several sites simultaneously. In other
words, we need to consider spheres of smaller radius
R
50
100
μ
m and only count the volume of overlap. Based on these
considerations we estimate the fraction of resolvable neurons to
be 13% (26%) for
R
=50
μ
m (75
μ
m), respectively. These
numbers critically depend on the high degree of sphere overlap
achieved by packing electrodes very densely (20
24
μ
m) along
the shanks, which was only possible through the use of nano-
scale interconnects.
In summary, we describe the design, construction, and valida-
tion of a con
fi
gurable system for dense 3-D electrophysiology.
While the in vivo experiments presented above establish the
proper operation of the system, signi
fi
cant additional work
remains to fully characterize the quality of spike sorting, the
yield of identi
fi
able single units, the merit of current source
density (CSD) estimates, and the long-term performance of the
system under chronic conditions. These are all areas of signif-
icant theoretical and experimental interest, and the technology
presented here will likely accelerate progress in these domains.
For example, the highly spatially resolved electrophysiological
recordings can be leveraged to analyze and improve spike
sorting and CSD estimation procedures. In particular, the dense
data can
fi
rst be spatially subsampled to mimic common
recording con
fi
gurations, and then the algorithm performance
on the coarsened data can be evaluated against the full set of
observations. Such cross-validation approaches can be used to
tune algorithm parameters and place bounds on error rates,
thereby mitigating the scarcity of ground truth data.
The technology itself can be further improved by scaling up
the number of recording sites by up to an order of magnitude,
while maintaining the modular architecture and small displace-
ment volume of the arrays. This would require the use of mul-
tiple interconnect layers on the shanks, denser interface matrices
at the probe base, multilayer
fl
exible cables, and higher channel
count signal conditioning ASICs. Since all of these requirements
e
ff
ectively bring traces closer together, the parasitic capacitance
budget is likely to be exhausted
fi
rst.
16
Beyond this point, active
components will have to be cointegrated closer to the recording
sites, in turn presenting the challenge of creating high-density,
yet low-power, active recording probes.
11
ASSOCIATED CONTENT
*
S
Supporting Information
The Supporting Information is available free of charge on the
ACS Publications website
at DOI:
10.1021/acs.nano-
lett.6b02673
.
Figure 5.
In vivo electrophysiological recordings using a 3-D array with 1024 electrodes. (a) Broadband signal (0.1 Hz
7.5 kHz) from the
hippocampus of an awake mouse from 4 of the 16 identical shanks (left) comprising the 3-D array. Each column displays 2 s of data from a single
shank with traces ordered by the depth of the corresponding microelectrode site. Notice that the spatiotemporal structure in the signal re
fl
ects the
anatomy and activation of the underlying circuit (cell layer marked by gray line). High-amplitude spikes are clearly visible on sites close to the cel
l
layer (pink, orange, and blue insets). (b) Similar spiking activity is seen throughout the array for sites near the pyramidal cell layer. (c) Histolog
ical
section showing the location of the shanks in panel a.
Nano Letters
Letter
DOI:
10.1021/acs.nanolett.6b02673
NanoLett.
2016, 16, 6857
6862
6861
Neural probe fabrication; h
eat dissipation analysis;
surgical procedures and in vivo recording methods;
histology; recording analysis and validation (
PDF
)
AUTHOR INFORMATION
Corresponding Authors
*
E-mail:
thanos@caltech.edu
.
*
E-mail:
roukes@caltech.edu
.
Author Contributions
G.R. and E.V.L. contributed equally to this work. System
architecture and speci
fi
cation: G.R., E.V.L., M.L.R., A.G.S.;
component design: G.R., E.V.L.; component mask layout: G.R.,
E.V.L.; component fabrication prototyping: G.R., D.C.; com-
ponent foundry fabrication: G.R., E.V.L., M.L.R., A.G.S.; system
packaging: G.R., E.V.L.; system software development: E.V.L.;
system benchtop evaluation: G.R., E.V.L.; system in vivo eval-
uation: G.R., E.V.L., A.G.S.; manuscript preparation and
fi
gure
creation: G.R., E.V.L. with input from A.G.S., M.L.R., D.C.
Notes
The authors declare no competing
fi
nancial interest.
ACKNOWLEDGMENTS
Foundry fabrication was carried out at CEA/LETI, Grenoble
under the aegis of the Alliance for Nanosystems VLSI; we
especially thank Denis Renaud, Eric Rouchouze, and Hughes
Metras for their help. We thank Jennifer Mok for histological
processing of brains from the in vivo experiments. This work
was supported by the Mathers Foundation, the Beckman
Institute at Caltech, the Moore Foundation, and the NIH
(1DP1OD008255/5DP1MH099907).
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