Materials Sciences and Applications, 2013, 4, 483-495 Published Online August 2013 (
Copyright © 2013 SciRes. MSA
Design and Implementation Challenges of Microelectrode
Arrays: A Review
Bahareh Ghane-Motlagh, Mohamad Sawan
Polystim Neurotechnologies Laboratory, Department of Electrical Engineering, Polytechnique Montreal, Montreal, Canada.
Received June 5th, 2013; revised July 4th, 2013; accepted July 16th, 2013
Copyright © 2013 Bahareh Ghane-Motlagh, Mohamad Sawan. This is an open access article distributed under the Creative Com-
mons Attribution License, which permits unrestricted use, distribution, and reproduction in any medium, provided the original work
is properly cited.
The emerging field of neuroprosthetics is focused on design and implementation of neural prostheses to restore some of
the lost neural functions. Remarkable progress has been reported at most bioelectronic levels—particularly the various
brain-machine interfaces (BMIs)—but the electrode-tissue contacts (ETCs) remain one of the major obstacles. The suc-
cess of these BMIs relies on electrodes which are in contact with the neural tissue. Biological response to chronic im-
plantation of Microelectrode arrays (MEAs) is an essential factor in determining a successful electrode design. By al-
tering the material compositions and geometries of the arrays, fabrication techniques of MEAs insuring these ETCs try
to obtain consistent recording signals from small groups of neurons without losing microstimulation capabilities, while
maintaining low-impedance pathways for charge injection, high-charge transfer, and high-spatial resolution in recent
years. So far, none of these attempts have led to a major breakthrough. Clearly, much work still needs to be done to
accept a standard model of MEAs for clinical purposes. In this paper, we review different microfabrication techniques
of MEAs with their advantages and drawbacks, and comment on various coating materials to enhance electrode per-
formance. Then, we propose high-density, three-dimensional (3D), silicon-based MEAs using micromachining methods.
The geometries that will be used include arrays of penetrating variable-height probes.
Keywords: Microelectrode Arrays; Electrodes-Tissues Contacts; Microelectrode Impedance; Neural Prosthesis
1. Introduction
Nervous system disorders reduce the quality of life of
afflicted persons—and are the most difficult to treat of all
parts of the human body. Most of these impairments re-
sult from lost localized sensory or motor dysfunction.
The evolving field of neuroprosthetics focuses on design
and implementation of neural prostheses such as visual
and cochlear implants to restore lost neural function by
recording signals from active neurons and subsequent
selective electrical stimulation of a population of neurons
[1]. Images or sounds taken from cameras or micro-
phones are coded to electrical impulses to stimulate reti-
nal tissues for the blind or auditory nerves for the deaf. In
another type of neuroprosthesis, electrical signals are
sent to a targeted portion of a patient’s brain suffering
from epilepsy, Parkinson’s disease, or depression to de-
crease the effects of the condition.
Motor neuroprostheses control external devices such
as prosthetic limbs with signals from the central or pe-
ripheral nervous systems. The ones that interface with the
central nervous system are called brain-machine inter-
faces (BMIs) [2]. The most critical component in this
function restoration is a microelectrode array (MEA)
implanted in neural tissue, which acts as an elec-
trode-neural tissues interface. In addition to injecting a
charge during stimulation, the ultimate role of MEAs is
to provide precise measurements of neural activity [3].
The field of implantable BMI based on MEAs is an
emerging research activity. Remarkable progress has
been reported at most bioelectronic levels, but the elec-
trode-tissue contacts (ETCs) remain one of the major
obstacles. Contacts achieved using MEAs do not comply
with the remaining parts of these BMIs due to the bio-
logical response to chronic implantation and to the elec-
tronic properties of MEAs. The success of these BMIs
relies on electrodes which are in contact with the neural
tissue. However, design and fabrication of a convenient
interface with selectivity, good electric characteristics,
sensitivity, biocompatibility, and long-term chemical and
recording stability remain a tough challenge [4]. For
Design and Implementation Challenges of Microelectrode Arrays: A Review
Copyright © 2013 SciRes. MSA
electrical stimulation and recording, we need electrodes
with high selectivity, low impedance, high-density, and
multi-dimensional geometry. Electrodes with high selec-
tivity and low impedance are preferred to stimulate spe-
cific nerve cells and to achieve a high signal-to-noise
ratio (SNR). For higher selectivity, however, a smaller
electrode size is required; this results in higher imped-
ance and a lower SNR [5]. To address this issues, we
should consider the entire combination of main as-
pects—design, materials, and fabrication techniques.
To date, implantable MEAs have been fabricated by
three common techniques: microwire, micromachined,
and flexible arrays [6]. Microwires, made from tungsten
or stainless steel, were the first implantable electrodes to
record chronically from the brain. This technique is used
to focus on the individual neuron [7]. One particular ad-
vantage of microwires is that they can be applied to ac-
cess deep brain structures. By using the arrays of mi-
crowires, simultaneous recording at the level of neuronal
populations is possible. One drawback of these elec-
trodes is that the accurate location of the electrode tips
relative to each other is not controllable because of the
wire bending during implantation. Micromachined elec-
trode arrays can be silicon or metal based. The sili-
con-based electrodes include two specific models. In the
first model, microelectrode sites are patterned on each
shank of silicon micro-machined substrate. This tech-
nique provides a higher density of sensors, which reduces
tissue displacement compared to microwires and contains
active electronics integrated into the arrays [8]. The sec-
ond model, however, includes sharpened silicon needles
electrically isolated from each other. The tips are coated
with platinum or iridium oxide; the rest are insulated
with polyamide. These electrodes are designed to be im-
planted in the cerebral cortex or peripheral nerves [9].
Another type of micromachined electrode is a met-
al-based array created by an electrical discharge machin-
ing (EDM) technique associated with electrochemical
steps. In this method, electrodes are made by microma-
chining a piece of stainless steel. After forming the elec-
trodes, platinum is deposited on the surface of the tips of
the electrodes which are electrically insulated from each
other by applying epoxy between them [10,11].
A flexible multi-electrode array is another type of
electrode. In this technique, the metal-based electrodes
are covered by two layers of polyimide, parylene, or
benzocycobutene without a silicon structure. Flexible
arrays provide an advantage over rigid electrode arrays
because of the closer mechanical match with brain tissue.
The mechanical mismatch between the rigid probes and
soft brain tissue can cause inflammation in the places of
implantation [12]. The flexible nature of these electrodes,
however, involves some difficulties during insertion.
Polymer-based implantable electrodes are a new ap-
proach to modify traditional electrode recording sites to
improve long-term performance [5].
In recent years, MEA techniques have been developed
to have a long-term and stable interface with the brain.
Different research projects are conducted to obtain con-
sistent recording signals from small groups of neurons
without losing microstimulation capabilities, while main-
taining low-impedance pathways for charge injection,
high-charge transfer, and high-spatial resolution by al-
tering the material compositions and geometries of the
arrays [13]. So far, none of these attempts have led to a
major breakthrough. One of the main challenges in fab-
ricating MEAs is finding appropriate coating materials
which will be biocompatible and improve neural re-
cordings. Conducting polymers and carbon nanotubes
(CNTs) have attracted much interest as suitable materials
for coating the electrodes. They ameliorate neural re-
cordings by decreasing electrode impedance and in-
creasing charge transfer but they cannot eliminate glial
encapsulation [4]. The molecular/cellular biology ap-
proach attempts to minimize the immune response to
implanted electrodes by using bioactive molecules. Such
an approach is focused on coating electrodes with bioac-
tive molecules [14]. To achieve these goals, we propose
3D, high-density, silicon-based MEAs using microma-
chining techniques. The geometries that will be used in-
clude arrays of penetrating variable-height probes. These
arrays can selectively communicate with large numbers
of individual neurons. To this purpose, the tips of sharp-
ened silicon needles will be coated with conducting
polymers using electrochemical techniques to enhance
both recording and electrical stimulation by decreasing
the electrode impedance and increasing charge transfer.
In this review, we describe the technology and design of
different types of microelectrode arrays with their ad-
vantages and drawbacks in Section 2. In Section 3 we
focus on about electrode-tissue interface, tissue reaction
to implantable electrodes which is referred to the bio-
compatibility of the implant, and some of the potential
strategies to improve the biocompatibility of implanted
cortical electrodes. Then, we propose our MEAs using
micromachining method.
2. Types of Electrodes
The design and implementation of appropriate implan-
table neural interfaces for electrical recording and efforts
to understand the mechanism of the nervous system at
the cellular level was begun in the last century; however,
research on using electrical stimulation to restore move-
ment of limbs was started more than two centuries ago
[15,16]. Neural interfaces are used to understand physio-
logical processes at the cellular level and to restore sen-
sory or motor function in the nervous system. Significant
progress has been achieved in the first part but there has
Design and Implementation Challenges of Microelectrode Arrays: A Review
Copyright © 2013 SciRes. MSA
been a lack of technological advances at restoration
nervous system dysfunction.
2.1. Microwire Arrays
The first implantable electrodes as an electronic interface
were made from sharpened wire metal in the 1950s.
Their narrow structures let them be placed very close to
single neurons in vivo, causing minimal damage in tissue.
They are insulated, except for the tips, to record the
highly localized extracellular potentials from the neurons
closest to the tips and also to inject the localized stimula-
tion current to excite the nearest neurons [17]. Metal
electrodes are fabricated from different types of materials
such as stainless steel, tungsten, platinum, iridium, or
gold. Strumwasser used 80-µm diameter stainless steel
wires to record discharges from single neurons of unan-
esthetized animals for a week or more [7]. Electrolyti-
cally sharpened tungsten wire electrodes with platinised
tips were used for extracellular single recording by Rob-
inson in 1968. The platinising process—a spongy deposit
of platinum black on the tips—enhances the effective
surface area and causes metal microelectrodes to achieve
their priority over glass pipettes [18]. Since stainless steel
is fragile near the tips, tungsten was replaced due to the
stiffness and rugged structure, and to provide very stable
recordings; however, tungsten is very noisy at low fre-
quencies. Signals of the mammalian nervous system
compared to other biological signals are fast and, by us-
ing a high-pass filter to remove noise, slow signals can
be lost; thus, tungsten is not a good choice for recording
these kinds of signals [19]. Metal electrodes have less
SNR because of lower impedance for the frequency
range of spike signals. A platinum electrode plated with
platinum black gives stable recordings, high SNR, and
creates a porous low-impedance structure, but it is me-
chanically fragile [20]. Iridium metal wire is extremely
stiff, highly resistance to corrosion, and its surface is
electrochemically activated, which causes it to increase
the maximum charge density [17].
Microwires remain in use today and fabrication me-
thods have not changed basically. They give long-lasting
individual neurons recording, sometimes more than one
year, so they allow neuroscientists to focus on individual
neurons. Using arrays of microwires allows multi-neural
recordings from neuronal populations [21]. Therefore,
simultaneous recordings from a large number of neurons
will be provided. Figure 1 shows a typical microwire
arrays fabricated by Plexon.
In all the above-cited cases, the final contact between
electrodes and brain tissue is a metal. Recently, CNTs
and polymers have been used to coat the tips of the metal
wire electrodes. In 2008, a group of researchers in Texas
coated the tips of conventional tungsten and stainless
steel wire electrodes with a combination of CNTs and
(a) (b)
Figure 1. Implantable plexon microwires: (a) Microwire ar-
rays attached to the high-density connector; (b) High-den-
sity microwire arrays [18].
polymer for the first time. They deposited CNTs on in-
dium-tin oxide MEAs [22]. Measurements showed that
the impedance decreased from 940 k to 38 k at a fre-
quency of 1 kHz, and showed ~40-fold increase in charge
transfer before and after coating (Figure 2).
CNT coatings can be applied to different materials and
geometries. In Figure 3 sharpened tungsten and stainless
steel wire electrodes were coated with CNTs.
The electrical characterization of modified tungsten
and stainless steel electrodes by electrochemical deposi-
tion of a covalent attachment scheme (CNT/CoV), and
CNT/polypyrrole (CNT/Ppy) composite are shown in
Table 1. Multi-walled carbon nanotubes (MWCNTs) add
to Ppy throught two different methods: layering and co-
deposition. The advantage of this combination is that the
MWCNTs are restricted within the polymer matrix, pre-
venting MWCNTs accumulation or mobilisation sur-
rended by cellular tissues [23]. CNT-coated electrodes
not only provide an appropriate substrate for neural
growth and function for at least three months but also
improve both the recording and stimulating characteris-
tics of neural electrodes [4].
One of the advantages of metal electrodes is access to
deep brain structures to treat disorders such as epilepsy,
Parkinson’s disease, and paralysis. By using arrays of
microwires, simultaneous recording at the level of neural
populations is possible [21]. One particular disadvantage
of microwires is bending of the wires during implanta-
tion, so the accurate location of the electrode tips relative
to each other is not controllable [6]. There is another is-
sue with metal electrodes—they are not always compati-
ble with silicon-based integrated circuits.
2.2. Micromachined Arrays
2.2.1. Silicon-Based Multi-Shaft Array
With the invention of the lithography technique in 1959,
which provided small size and accurate dimensional con-
trol for integrated circuit fabrication and use of this
process for fabricating extracellular MEAs, a new way
for simultaneously recording from many neurons, or
Design and Implementation Challenges of Microelectrode Arrays: A Review
Copyright © 2013 SciRes. MSA
(a) (b)
Figure 2. Characteristics of CNT-coated MEAs: (a) An im-
pedance spectroscopy scan; (b) A cyclic voltammetry scan
(a) (b)
Figure 3. Metal electrodes coated with CNTs: (a) CNTs
covalently attached to the sharp tungsten electrode, (b)
Parylene insulation exposed stainless steel electrode re-
moved by UV laser, polymerized with a mixture of CNTs/
Polypyrrole composite [4].
stimulate them, was provided [24,25]. Use of the litho-
graphic technique and silicon etching technology to cre-
ate high-density electrode arrays for single-unit recording
began at Stanford University in 1966. By the late 1960s,
these new technologies were used by Brindley to fabri-
cate visual prostheses and they were developed as a cor-
tical prosthesis for the blind at the University of Utah
[26,27]. Silicon-based electrodes intended for both
long-term monitoring of neural activity and injection of
stimuli into neural tissues at the cellular level are com-
patible with CMOS technology, which uses planar pho-
tolithographic techniques on silicon wafers. Despite the
problem that microwire tips may move apart during im-
plantation, the spatial relation between electrodes in mi-
crofabricated arrays is determined during the microfab-
rication process and remains fixed. One of the first sili-
con-based MEAs was made at Michigan University dur-
ing the 1970s. In this design, electrode sites are placed on
each shank. This kind of electrode can easily merge with
on-chip circuitry, signal processing, and wireless inter-
faces. Figures 4 and 5 show the basic structure of a mi-
cro-machined microelectrode probe for recording or
stimulation in the central nervous system and a four-pe-
Table 1. Electrical properties of nanotube-modified electo-
des [4].
Normalized impedance
Normalized capacitance
CNT/CoV0.075 38
CNT/Ppy 0.77 755
Figure 4. Typical Michigan probe: Basic structure of a mi-
cromachined microelectrode probe [29].
netrating-shank 16-channel probe with four uniformly
recording sites, respectively [28,29]. The silicon sub-
strate shank is 15 µm thick, 3 mm long, and 90 µm wide
at the base, narrowing to 20 µm at the tip. The surface
area of the recording sites is between 100 to 400 µm2.
The high-density, 3D electrode arrays are inserted in the
tissue to record electrical activity or inject electrical sig-
nals to stimulate neurons. The substrate of the probe
(shank) should be biocompatible, strong, and small
enough to penetrate to the pia arachnoid membrane over
the brain. Silicon is strong enough and allows integrated
electronics to be formed directly in the probe substrate
The recording sites are typically made of iridium (Ir)
or gold. These recording sites can be modified by coating
with conductive polymers such as Ppy or poly (3, or 4-
ethylenedioxythiophene-PEDOT) to improve long-term
performance and decrease recording site impedance. Ex-
perimental and theoretical characterization of these mi-
croelectrodes modified with conducting polymer nano-
tubes showed a significant impedance decrease and a
charge injection capacity increase by two and three or-
ders of magnitude, respectively. The impedance spec-
troscopy results of PEDOT nanotube-coated electrode
sites with applied charge inject density of 1.4 C/cm2 are
shown in Figure 6.
The charge capacity of bare gold, Ppy, and PEDOT
electrodes are shown in Table 2 [31].
According to the CNT-related results, which included
Design and Implementation Challenges of Microelectrode Arrays: A Review
Copyright © 2013 SciRes. MSA
Figure 5. Typical Michigan probe: General schematic of the
four-shank electrodes forming16-channel probe [28].
Figure 6. Measured and calculated data of PEDOT nano-
tubes impedance with applied charge density of 1.4 C/cm2:
(a) Bode plot; (b) Nyquist plot [31].
decreasing the impedance and noise, the recording sites
were modified by CNTs coating [32]. The results showed
that CNTs incorporation by increasing the effective sur-
face area could decrease site-tissue impedance [33].
Novel, multi-sided parylene-based MEAs with elec-
trode sites in the front-side, back-side, and on the edge of
the device was fabricated for neural recording. This me-
thod can increase the function of each shank [34].
One of the advantages of these microelectrodes is pro-
viding a high density of sensors with predetermined loca-
tions and high spatial resolution. There are some issues
with this kind of electrode array: the shank—which in-
cludes recording sites—causes more tissue displacement
and may damage some neurons during insertion; also,
dura mater should be partially removed for inserting the
electrode arrays (Figure 7). The voluminous electronic
structure, which assembles the electrode shanks, causes
more damage in the brain tissue during insertion to the
2.2.2. Silicon-Based Multi-Needle Array
Another type of silicon-based microelectrodes is Utah
arrays which consist of conductive, sharpened silicon
needles electrically isolated from each other. The maxi-
mum length of the needles is 1.5 mm. The tips of the
needles are coated with platinum or iridium oxide; the
rest are insulated with biocompatible polymers such as
polyimide or parylene-C. The architecture of these elec-
Table 2. Charge capacity of bare gold, Ppy, and PEDOT
electrodes with applied charge density of 1.4 C/cm2 [31].
Electrodes Charge capacity (µC)
Gold 0.001 ± 104
Ppy 2.3 ± 0.5
PEDOT 4.9 ± 0.6
(a) (b)
Figure 7. The Michigan cortical implant: (a) A 3D view of
the microsystem which includes implanted electrodes with
ribbon cables connecting them to an electronic device, in-
cluding circuitry for amplification, spike detection, encod-
ing, and wireless transmission of power and bidirectional
data; (b) Four 64-site probes assembled into a 3D structure
trodes enables single-unit recording with high spatial
resolution, as well as exciting the neurons by electrical
stimulation. An array of electrodes with small surface
areas and which penetrate to the cortex stimulate more
localized regions of tissue [35]. Such microelectrode
arrays can also be safely inserted into the brain [36]. To
use microelectrode arrays for recording and stimulation,
electrical and mechanical characteristics are important.
Geometrical architecture is the parameter, which defines
those characteristics in electrode arrays. Due to the Utah
microelectrode arrays’ architecture, they have been used
extensively in neuroscience and clinical research. Figure
8 shows the two available architectures of Utah arrays,
which count 100 microelectrodes with platinum-coated
tips [37]. The Utah slanted electrode arrays provide 2D
recording data from the slanted plane of the brain
whereas the flat array provides 1D data. The issue with
this type of arrays is that even the slanted array is still 2D
instead of 3D. The latter microelectrodes are designed to
be implanted in the cerebral cortex or peripheral nerves
[1]. To facilitate charge transfer from the electrodes to
neural tissue, the tips of the electrodes are coated with
platinum (Pt), sputtered iridium oxide film (SIROF), ti-
tanium nitride (TiN), and soft conductive polymers such
as Ppy and PEDOT. Pt is a noble metal and injects
charge by both Faradaic reactions and double layer
Pt charge inject capacity using 0.2 ms pulses is 50 -
150 µC/cm2. Iridium oxide (IrOx) is using to coat the tips
due to its ability to inject charge, its resistance to corro-
Design and Implementation Challenges of Microelectrode Arrays: A Review
Copyright © 2013 SciRes. MSA
(a) (b)
Figure 8. Utah silicon-based 100 microelectrode arrays: (a)
2D flat electrode array; (b) 2D slanted electrodes array [1,
sion, and high conductivity. IrOx is a highly conductive
oxide with bulk resistivity between 30 to 100 µ·cm.
Using 0.2 ms pulses, the activated IrOx charge inject ca-
pacity will be 1 mC/cm2 [38]. TiN is another material
found to be useful due to chemical stability and biocom-
patibility. Using 0.5 ms pulses provides 0.9 mC/cm2
charge injection capacity [39]. Table 3 shows the charge
injection capacity of above materials. The protein ad-
sorption of materials are being used for recording sites of
implantable MEAs is another important factor to deter-
mine the stability of the implant. The absorbed protein
layers on gold are thinner than other materials such as
platinum and iridium. However, indium tin oxide might
be a good alternative for coating [40].
Conductive polymers are now considered promising
materials for chronic stimulation and recording due to
their biocompatibility and large surface area which acts
as a main factor to reduce the impedance of the elec-
trodes. PEDOT nanotubes have a low impedance of 4 k
at 1 kHz, which increases the charge injection at the elec-
trode-tissue interface.
Figures 9(a) and (b) shows the stability of the SIROF-
and Pt-coated electrodes in vitro over 90 days and a
typical Bode plot as a function of frequency. The median
impedance of Pt- and SIROF-coated electrodes were 125
± 0.25 k and 6.7 ± 0.2 k at 1 kHz, respectively. The
impedance of uncoated silicon tips was 975 ± 15 k at 1
kHz. A typical Bode plot presents the impedance of the
Pt- and SIROF-coated Utah electrodes versus frequency.
According to the plot, the impedance of the SIROF is
lower than the Pt electrodes at all frequencies except at
105 Hz, which is a high frequency [41].
2.2.3. Metal-Based Microelectrode Array
These microelectrodes arrays were fabricated by using a
wire EDM technique associated with electrochemical
steps [10]. The electrodes were made from stainless steel
or titanium. After chemical etching to smooth the elec-
trode surface, platinum was electro-deposited at the tips
to facilitate charge transfer from electrode to neural tis-
sue. To electrically insulate the electrodes parylene-C
Table 3. Charge injection capacity of coating materials [41].
Electrodes Pulse width (ms) Charge capacity (µC/cm2)
Pt 0.2 50 - 150
IrOx 0.2 1000
TiN 0.5 900
was chosen as an insulator due to its biocompatibility.
Electrodes were insulated from each other by epoxy [11].
Figure 10 shows a general view of these fabricated elec-
trodes. This microelectrode array was developed to be
assembled with integrated circuits on a thin substrate.
[42]. The main drawback of these electrodes is that they
are not compatible with integrated circuit manufacturing
techniques. The impedance of the electrodes was ob-
tained from the impedance curves for each electrode us-
ing an impedance analyser. Measurements were per-
formed at 25˚C in 0.9% saline; 43 out of 64 electrodes
had an average impedance of 78 k, and 21 of electrodes
had an average impedance of 5.7 M.
2.3. Flexible Array
A variety of MEAs has been developed for recording and
monitoring neural activities; however, most MEAs are
based on a rigid substrate and cause neural damage and
inflammation at the implant site for intra-cortical im-
plantation. Flexible arrays allow electrodes to fit on the
surface of the brain to record and monitor in a less inva-
sive way. Flexible, polyamide-based microelectrodes are
used for chronic neural recording because they provides
high mechanical flexibility, good biocompatibility, high
resistance to solvents, and are a strong addition to metal
oxides [43]. Different polymers such as polyimide and
parylene are used as a structural substrate. Parylene-ba-
sed MEAs are extremely flexible, providing highly con-
formal coverage of the muscle surface and a stable elec-
trical contact, which results in an improved SNR [44].
Recently, flexible CNT microelectrodes have not only
improved electrode impedance and charge-transfer ca-
pacity but also—when the CNTs are grown directly on
the polyimide substrate—their adhesion to the substrate
is improved [45].
Flexible parylene-based microelectrode arrays for re-
cordings are shown in Figure 11.
A new parylene-based multi-sided MEA with elec-
trode sites at the top-side, back-side, and edge has been
presented for neural recording and passive drug delivery.
This feature creates the smallest footprint (85 μm2) to
date of a functional recording electrode. Gold microelec-
trodes have an average impedance of 6.0 M at 1 kHz
which, by using PEDOT electropolymerization, reduce
the impedance to 210 k. In fact, by increasing the elec-
trochemical area, PEDOT polymerization decreases the
Design and Implementation Challenges of Microelectrode Arrays: A Review
Copyright © 2013 SciRes. MSA
(a) (b)
Figure 9. Stability of Pt- and SIROF-coated electrodes in vitro over 90 days; (b) The Bode plot presents the electrode imped-
ance versus frequency of Pt- and SIROF-coated Utah electrodes [41].
(a) (b)
Figure 10. Metallic-based microelectrode arrays using an EDM technique: (a) SEM image of a parylene-coated assembly,
consisting of platinum-coated electrode tips; (b) Stainless steel electrode arrays after electrochemical polishing [10,11].
(a) (b)
Figure 11. Flexible parylene-based electrodes: (a) Global
view of eight microelectrodes; (b) Photomicrograph picture
of microelectrodes with two different spaces between the
recording sites; (c) Photomicrograph picture of a single
recording site that includes a gold pad and a parylene-in-
sulated region (dark yellow) [44].
overall impedance. A different design of novel pary-
-lene-based microelectrodes for recording, passive drug
delivery, and cell delivery are shown in Figure 12 [34].
Table 4 shows the impedance of the gold and PE-
DOT-coated recording sites at the surface and the edge.
Flexible surface electrodes provoke less tissue damage
because the surgeon may avoid penetrating neural tissues;
however, surface electrodes have low sensitivity and
selectivity. In addition, because of its flexible nature,
there is difficulty inserting electrodes into the nervous
3. Electrode-Tissue Interface
Electrical coupling and biocompatibility are the most
important issues at electrode-tissue interfaces. While
electrical coupling is determined by the impedance at the
interface, biocompatibility is influenced by the materials
at the recording or stimulation site [46]. Low impedance
causes low noise and is obtained with a large surface area.
A new approach to this issue is using nanomaterials such
as semiconductor quantum dots [47], metallic nanoparti-
cles [48], metallic or semiconductor nanowires, nanos-
Design and Implementation Challenges of Microelectrode Arrays: A Review
Copyright © 2013 SciRes. MSA
Figure 12. Different designs of multi-sided microelectrode:
(a) Front-side, back-side, and edge designs of electrodes
with 45-micron spacing; (b) Open-channel design for pas-
sive drug delivery; (c) Hollow-channel neural probe for
cell-based therapies; (d) Ring, sieve, cantilever, reservoir,
edge, PEDOT-coated edge electrodes [34].
tructured conductive polymers [49], carbon nanofibers
[50], or carbon CNTs that are characterized by excellent
electrical [51], mechanical, and chemical [52] properties
for neural interfaces. Electronic properties of the nano-
materials are match with the charge transport require-
ments of cellular interfacing besides the mechanical, and
the chemical properties of nanomaterials are critical for
long-term implants. Nanomaterials play a significant role
in improving electrode domains by increasing the sen-
sitivity and selectivity, improving response time, and
making the electrodes biocompatible with the biological
environment. Due to their rich electronic properties, high
aspect ratio, excellent chemical stability, high mechani-
cal strength, and intrinsic large surface area, CNTs are
one of the best options for the active site of electrodes
Other promising materials to coat nerve electrodes are
polymers with properties of their modification with bio-
molecules to improve biocompatibility and create a bio-
logically active electrode-tissue interface. PEDOT is an
electrical conductive polymer and the most promising
material to coat electrodes [54,55].
Implanted electrodes exhibit a short, effective life-
time due to degradation of signal transmission at the
tissue-electrode interface [56]. Optimum electrodes,
Table 4. Electrode sites impedance [34].
Au site,
N = 48 PEDOT-PSS1 site, N = 33
Planar 17 × 17 µm5.4 ± 3.7 M 240 ± 110 k
Edge 17 × 5 × 2 µm6.0 ± 3.1 M 210 ± 64 k
1PSS: Polystyrene Sulfonate.
which implant in neural tissue, should cover two impor-
tant characteristics: long-term tissue compatibility and fast
response to signals [57]. Long-term tissue compatibility is
related to a tissue response to a foreign body. This re-
sponse results in glial activation. Microglial cells, which
integrate the intrinsic brain defence system, are very sen-
sitive to variations in extracellular ionic concentrations
and form an insulating interface that prevents signal
conduction or charge injection for long-term implanta-
tion [58]. Local activation of glial cells causes electrode
encapsulation due to chronic electrode implant and it
increases the electrical impedance of the recording site
[59]. Different methods are being used to prevent cellular
encapsulation and electrode failure. Surface modification
of electrodes can improve the interaction between neural
cells and implants. L1, a neural adhesion molecule, is
one of the coating materials that may increase the bio-
compatibility of neural probes and promote electro-de-
tissue interface.
4. Comparison of Microelectrode
Findings to date demonstrate different methods for re-
cording and stimulating neuron activities; however, all
these methods have some drawbacks. The tables (Tables
5-7) show a brief comparison of various techniques for
fabricating MEAs.
5. Design and Microfabrication of MEAs
As described above, various types of MEAs exist; these
are silicon-based, treated with dedicated electrochemical
steps; however, these types of microelectrode arrays,
fabricated by micromachining techniques, still lack the
required 3D, high-density and low-impedance, necessary
for ETCs and required by implantable stimulators. By
focusing on these aspects the ultimate purpose of intro-
ducing new microelectrodes is to elaborate effective
MEAs with selectivity, sensitivity, biocompatibility,
good electric characteristics, and long-term chemical and
recording stability. To achieve this goal, a silicon-based
micromachined method was chosen to make the new
microelectrodes due to compatibility with integrated cir-
cuit manufacturing techniques and a high degree of se-
lectivity in stimulation and recording. Front-side diced
electrodes were etched after backside dicing, glassing,
and metallization. As a result, we obtained electrode ar-
Design and Implementation Challenges of Microelectrode Arrays: A Review
Copyright © 2013 SciRes. MSA
Table 5. Advantages and drawbacks of MEA fabrication methods [6,9,24,60].
Method Advantages Drawbacks
Microwire arrays
Long-lasting and single-neuron recording for more than a year
in some cases.
Access to deep brain structures.
Multi-neuronal recordings for simultaneous recording at the level
of neural populations.
Electrode tips location cannot be controlled (to
bend during implantation).
No embedded microfluidic channel.
Incompatible with electronic integrated circuits
which are silicon based.
Spatial relation between the electrode sites remains fixed.
Reproducible with high spatial resolution (thin-film structures).
Facilitates current source density analysis because of linear
arrangement of recording sites along a single shank.
More tissue displacement because of the shank.
Dura mater should be removed for inserting the
More damages in the brain tissue due to large
electronic structure.
Rigid and may fracture.
High degree of selectivity in stimulation and recording.
Ease of implantation.
Compatible with IC manufacturing technique.
Rigid and may fracture.
May damage neural tissues during insertion.
Stimulate more localized regions of tissue.
Easy to insert (little pin units). Incompatible with silicon-based IC technology.
Flexible arrays
Good electrical insulation and adhesion properties.
Conformal coverage with the surfaces.
Flexibility reduces the chronic tissue inflammation response.
Flexibility nature implies difficulty in insertion.
Table 6. Characterization of MEA technologies [31,41].
MEAs MEAs sitesMean charge density (mC/cm2) Mean Impedance-Z1kHz (k)
SIROF 96 2 6.1 ± 0.2
Utah array Pt 96 0.3 125 ± 0.25
Polypyrrole 19.5 ± 2.1 184 ± 5.3
Michigan array PEDOT 64 2.5 ± 1.4 392 ± 6.2
TDT1 Microwires 48 5.10 ± 0.40 19.9 ± 0.82
1TDT = Tucker Davis Technologies.
Table 7. Comparison of microelectrode structure [43,61-63].
MEAs Site/probe geometry Substrate material Recording site material
Diameter: 50 μm
Electrode spacing: 400 µm
Tip angle: 45˚
Polyimide insulation/Parylene-C insulation Tungsten, Stainless steel, Pt,
Utah array
Shank length: 1.5 mm
Base: 1.6 mm2
Electrode spacing: 400 µm
Conical recording site
Boron doped Silicon/Parylene-C insulation Pt, IrOx, 1SIROF
Shank length: 3mm
Probe thickness: 15 μm
Probe width: 33 - 55 µm
Electrode spacing: 100 µm
Silicon/Silicon dioxide/
nitride insulation/Parylene Gold, Ir, 2PEDOT, CNTs
Flexible arrays Shank thickness: 15 µm Polyimide/ Parylene/ Benzocyclobutene Gold, 3Ppy, PEDOT
1SIROF: Sputtered Iridium Oxide Films; 2PEDOT: poly (3, or 4-ethylenedioxythiophene); 3Ppy: Polypyrrole.
rays with variable heights of 1.45, 1.55, and 1.65 mm.
The thickness of the electrodes was 200 µm at the base
and 5 µm at the tip with 98 µm spacing. Figures 13(a)
and (b) shows the preliminary result of fabricated MEAs
and depict the desired 3D electrode array. We are fabri-
cating 3D high-density silicon-based MEAs. Such 3D
silicon-based microelectrodes in an array will allow not
only recording from different depths of the brain but
cause less tissue damage during insertion.
To enhance the high-selective, low-impedance, and
high signal-to-noise ratio necessary for ETCs, fabricated
silicon probes capable of simultaneously recording and
stimulating from neuronal populations will be coated
with various conductive polymers and CNTs.
6. Discussion
The emerging field of neuroprostheses that incorporates
BMIs is being developed to solve nervous system disor-
ders that people suffering from. To date the technology
for creating MEAs capable of restoring lost neural func-
tion is still in progress. Clearly, much work still needs to
be done to accept a standard model of MEAs for clinical
purposes. Electrical and mechanical characteristics of
Design and Implementation Challenges of Microelectrode Arrays: A Review
Copyright © 2013 SciRes. MSA
(a) (b)
Figure 13. Microelectrode arrays: (a) SEM image of the MEAs; (b) Schematic of proposed 3D MEAs.
electrode arrays are important for stimulation or recor-
ding. Geometrical characteristics of the electrodes in an
array will affect the electrical and mechanical properties.
Mechanical conformity minimizes tissue damage during
insertion. 3D electrodes with variable heights of the elec-
trodes in an array will provide more contacts between the
electrodes and targeted neural tissue. More importantly,
the novel geometry will allow not only recording from
different depths of the brain but will cause less tissue
damage during insertion. To achieve high-density elec-
trode arrays in this proposed structure, the distance be-
tween electrodes is decreased. Therefore, the density of
pins will be higher, which will allow us to selectively
stimulate in various sites in the 3D area of the MEAs. To
achieve electrodes with higher selectivity and less tissue
damage during penetration into the tissue, thinner and
sharper electrodes are desired. Besides, small tip expo-
sures increase the electrode impedance. To improve neu-
ronal signals, the impedance of the electrode should be
low. To achieve this goal, the tips of the electrodes will
be coated with conductive polymers and CNTs to im-
prove the recording and stimulation characteristics of the
neural electrodes. The coating enhances electrical proper-
ties by decreasing electrode impedance and increasing the
charge transfer.
7. Acknowledgements
We gratefully acknowledge the financial support from the
Natural Sciences and Engineering Research Council of
Canada (NSERC) and the Canada Research Chair in
smart medical devices, and design tools and fabrication
resources from CMC Microsystems.
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