REVIEW
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Electrocorticogram (ECoG): Engineering Approaches and
Clinical Challenges for Translational Medicine
Hyunmin Moon, Jii Kwon, Jonghee Eun, Chun Kee Chung, June Sic Kim,*
Namsun Chou,* and Sohee Kim*
subdurally. To detect such ECoG signals,
the most critical technological component is an electrode which must be in
contact with the cortical surface. Unlike intracortical electrodes, ECoG electrodes are minimally invasive since they
only touch the cortical surface and do
not penetrate the cortical tissue. Hence,
they generally cannot detect action potentials generated by individual neurons
but can detect the superposition of action potentials generated by multiple
neurons as a group, or even the superposition of subthreshold potentials.
The cortex is where motor functions
are commanded, sensory information
is received and processed, and higherlevel activities such as the association of
multimodal sensory information, judgment, and cognitive processing occur.
ECoG signals are generally regarded as a
reflection of the electrophysiological activities of multiple neurons simultaneously. Thus, ECoG has become a key tool for
detecting brain activity with higher-quality signals and improved stability and accuracy compared to noninvasive electroencephalography (EEG), but with less invasiveness compared to intracortical recording. In particular, ECoG electrodes are expected
Electrocorticogram (ECoG) is an electrophysiological signal that results from
the summation of neuronal activity near the cortical surface. To record ECoG
signals, the scalp and skull are surgically opened and electrodes are placed on
the cortical surface, either epidurally or subdurally. Owing to its improved
spatiotemporal resolution and signal quality compared with
electroencephalography, it is widely used to diagnose and treat neurological
disorders in clinical settings for several decades, despite the invasiveness of
ECoG. Recently, ECoG is applied in research to explore brain functions and
connectivity, brain-computer interfaces, and brain-machine interfaces. In
addition to the need for ECoG in neuroscience research, ECoG devices have
advanced in terms of materials, fabrication, and function to overcome the
limitations of commercially available ECoG arrays. Here, the conventional use
of ECoG in clinical medicine, the new applications of ECoG in basic
neuroscience research, and the future challenges in translating recent
developments in ECoG devices for clinical use are described.
1. Introduction
An electrocorticogram (ECoG) is an electrical signal resulting
from the summation of electrophysiological activities from multiple neurons residing near the cortical surface. ECoG signals
can be detected from the cortical surface, either epidurally or
H. Moon, S. Kim
Department of Robotics and Mechatronics Engineering
Daegu Gyeongbuk Institute of Science and Technology (DGIST)
Daegu 42988, Republic of Korea
E-mail: soheekim@dgist.re.kr
H. Moon
Department of Mechanical Engineering
Massachusetts Institute of Technology
Cambridge, MA 02139, USA
J. Kwon, C. K. Chung
Department of Brain & Cognitive Sciences
Seoul National University
Seoul 08826, Republic of Korea
The ORCID identification number(s) for the author(s) of this article
can be found under https://doi.org/10.1002/admt.202301692
© 2024 The Authors. Advanced Materials Technologies published by
Wiley-VCH GmbH. This is an open access article under the terms of the
Creative Commons Attribution License, which permits use, distribution
and reproduction in any medium, provided the original work is properly
cited.
J. Eun, N. Chou
Emotion, Cognition, & Behavior Research Group
Korea Brain Research Institute
Daegu 41062, Republic of Korea
E-mail: nschou@kbri.re.kr
C. K. Chung
Department of Neurosurgery
Seoul National University
Seoul 08826, Republic of Korea
J. S. Kim
Clinical Research Institute
Konkuk University Medical Center
Seoul 05030, Republic of Korea
E-mail: 20230375@kuh.ac.kr
S. Kim
Department of Electrical and Computer Engineering
University of California San Diego
La Jolla, CA 92093, USA
DOI: 10.1002/admt.202301692
Adv. Mater. Technol. 2024, 9, 2301692
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© 2024 The Authors. Advanced Materials Technologies published by Wiley-VCH GmbH
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to detect brain signals as accurately as intracortical recordings,[1,2]
as well as to stimulate a specific brain region to treat neurological
diseases.[3,4]
In this review, we introduce the ECoG electrodes and procedures currently used in clinical environments and review recent technological developments toward emerging applications
of ECoG beyond the current clinical use. Starting from the history of ECoG use in humans, the use of ECoG in clinics for diagnosis, treatment, analysis of brain functions, and brain-computer
interfaces (BCI) or brain-machine interfaces (BMI) is discussed.
Next, engineering efforts and approaches for developing novel
ECoG tools for utilizing ECoG in translational research are introduced. Such novel devices enable the detection of higherresolution signals using ECoG with additional functionalities.
The materials, fabrication methods, and additional functionalities used for micro-ECoG devices are highlighted. Finally, the existing challenges and directions for translating micro-ECoG electrodes into clinical applications are discussed.
2. Clinical ECoG
2.1. Brief History of Electrophysiology and ECoG
In 1875, Caton reported measuring electric currents from two
electrodes located on the surface of the gray matter and skull in
experiments using rabbits and monkeys.[5] Beck et al., in 1890,
showed the spontaneous and evoked electrical brain activity induced by sensory stimulation in dogs and rabbits.[6] The first EEG
recording of focal seizure-related cortical activity in dogs was reported by Cybulski and Macieszyna in 1914.[7–9] In 1929, Berger
reported the first noninvasive EEG recording in humans.[10] He
described changes in the alpha and beta waves depending on
whether the eyes were open or closed. In 1930, he performed the
first direct recording of the cortical surface and white matter in
humans.[11,12]
Subsequently, Berger first recorded epileptiform activities
from a human patient. In 1933, He reported in his report, “Über
das Elektrenkephalogramm des Menschen”, that he used EEG
to record high voltage waves three times per second during
seizures.[9] After Berger’s report, Gibbs et al. (1935) demonstrated EEG epileptic patterns associated with different types of
epilepsy-petit mal seizures and grand mal seizures;[13] later, they
showed the possibility of the diagnosis of epilepsy and localization of seizure origin using brain activities.[14] Foerster not only
confirmed Berger’s results about epileptiform activities but also
reported extremely low potential fluctuations in the region of
tumor using an electrode placed on the exposed surface of the
brain.[15]
In 1939, Penfield performed the first invasive monitoring using epidural ECoG with brain stimulation to localize the eloquent
and epileptogenic zones.[16–18] Hunter and Jasper suggested a
method for recording EEG activity and seizure patterns simultaneously using an optical system consisting of mirrors and a
camera.[19] In 1966, Goldensohn used closed-circuit television
(CCTV) combined with the recording of EEG signals for the first
time,[20] which became the current standard of invasive monitoring systems that simultaneously observe seizure patterns and
epileptiform activities.
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2.2. Recent ECoG Applications
The application of ECoG in human subjects has continued to
grow since the 2000s as depicted in Figure 1a. Previously, electrocorticography using strips or grids of electrodes implanted in
the subdural space was used to define the range of tumor resections for short-term intraoperative use[11,15] or to provide electrical stimulation for pain reduction in patients with intractable
neuropathic pain.[21] In recent clinical fields, ECoG electrodes are
generally implemented for invasive extra-operative monitoring
as an implanted use in patients with drug-resistant epilepsy and
are considered the gold standard for identifying precise seizure
onset zones for resective surgery.[22,23] In addition, direct cortical stimulation using ECoG electrodes can provide information
on the functional role of specific cortical sites and pathological
responses such as afterdischarge. Understanding the functional
role of a certain area is essential for identifying the eloquent areas that are crucial for recovery and quality of life after respective
surgery.[24–26]
In modern clinical settings, ECoG has been primarily used in
patients with epilepsy to detect epileptic zones before resective
surgery. In addition to epilepsy, ECoG has shown potential for
various clinical applications in the past, including tumor localization, pain reduction, Parkinson’s disease, obsessive-compulsive
disorder, and major depression (Figure 1b).[27] This growth is not
limited to clinical use but has expanded significantly into various
areas of neuroscience research. In certain cases, brain activities
from these patients have been read out for neuroscience studies to explore sleep, pain, cognition, and memory functions and
for BCI/BMI applications. Particularly, with the advancement of
BCI and BMI technologies, ECoG’s feasibility is increasing for
rehabilitation purposes in patients with conditions like lockedin syndrome, amyotrophic lateral sclerosis, and spinal cord
injury.[28]
2.2.1. Epilepsy Monitoring
The lifetime prevalence of epilepsy is known to be 7.60 per 1000
individuals worldwide.[29] Almost 1% of the world’s population
suffers from or has suffered from epilepsy. Approximately 25%
of these patients have drug-resistant epilepsy, which occurs when
an adequate trial of two anti-epileptic drugs (AEDs) fails to relieve
seizures.[30,31] For those drug-intractable patients, surgical treatment is performed to achieve seizure relief by localization and
removal of the seizure onset zone. ECoG electrodes are required
for precise surgical procedures.
Before implantation of the ECoG electrodes for invasive monitoring, a noninvasive diagnosis is performed preoperatively.[32]
To determine the laterality of the seizure onset zone, scalp
EEG recordings are first conducted, along with video recordings to confirm the seizure patterns. To identify anatomical abnormalities or functional impairments associated with seizures,
various imaging techniques are utilized, including magnetic
resonance imaging (MRI) to assess structural abnormalities,
positron emission tomography to evaluate glucose metabolism,
and single photon emission computed tomography to assess
regional cerebral blood flow. Sometimes biochemical abnormalities are evaluated with magnetic resonance spectroscopy.
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Figure 1. Trends of publications regarding electrocorticography. a) The number of publications searched using the term “electrocorticography” is represented by blue bars, and using the terms “electrocorticography” and “human” is represented by magenta bars. b) The number of publications on human
research is categorized into two fields of medicine (blue bar) and neuroscience (magenta bar). All data are based on Scopus.
Magnetoencephalography, another noninvasive method for detecting brain activity, is useful in diagnosing neocortical epilepsy,
where abnormalities are rarely found in structural images, and
can provide important information to localize eloquent areas.[33]
The Wada test (a sodium amobarbital intracarotid injection) and
neuropsychological tests can be used to evaluate lateralization
and quantitative functionalities such as attention, memory, language, and motor function.[34]
In combination with neuroimaging or other diagnostic modalities, invasive monitoring is performed before surgical removal
of the epileptogenic zone, distinguished from the eloquent area.
In invasive EEG monitoring, brain activity is recorded for a certain period by implanting electrodes, such as grid or strip ECoG
arrays or depth electrodes, onto the area presumed to be associated with a seizure origin using the aforementioned non-invasive
modalities. Seizure origins can be determined by detecting the
presence of spikes during interictal periods or determining the
area where the change in brain activity first appears during the
seizure (Figure 2a).
Through invasive ECoG studies in patients with epilepsy, it
is confirmed that there is a significant association between a
seizure-free outcome and resection areas showing a spreading
ictal signal within 3 s, pathological low-frequency waves, or frequent interictal spikes.[35] In human patients and animal studies,
high-frequency epileptic activity above 80 Hz has been recorded
in epileptic tissues.[36,37] It is reported that the extracted feature
from the phase locking value between the phase of high gamma
activities (80–150 Hz) and the phase of lower frequency rhythms
(4–30 Hz) is useful in finding the seizure onset zone.[38]
2.2.2. Study of Brain Mechanisms
In addition to epilepsy monitoring, ECoG has been used in clinical studies on sleep and consciousness. Slater et al. suggested a
novel measure of local sleep while awake by detecting changes
in local field potentials.[43] Alonso et al. proposed a method to
quantify changes in neural activity to characterize the loss of consciousness by observing the enhanced stability of neural dynamics in an anesthetized state.[44]
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In neuroscience, ECoG signals provide insights into the mechanisms of human cognitive processes. In particular, high gamma
activity is considered the neural population activity of neurons
located directly underneath the ECoG electrodes, which are induced by a specific task.[40,45–49] Therefore, such focal activities
represented by high gamma responses are considered to be significant and useful in cognitive neuroscience as well as in braincomputer interfacing.[50]
Crone et al. reported that an increase in gamma activity
(>30 Hz) is induced by auditory perception.[46] Additionally, during auditory processing, corticocortical connections were suggested by the causal relationships of high gamma activity between areas in the early auditory pathways.[48] Ryun et al. showed
that high gamma (50–140 Hz) power can be modulated by vibrotactile frequencies in the human primary and secondary somatosensory cortices (Figure 2b).[49] In the visual cortex, narrowband gamma activity was shown to differ depending on the
properties of the stimuli, such as white noise or grating, but
broadband gamma oscillations generally increased for all visual
stimuli.[51]
Many studies have been conducted not only on sensory tasks,
but also on cognitive experiments, such as memory, movement,
and language. Lee et al. found that the cortical network in
the alpha band (8–13 Hz) between the medial temporal and
parietal lobes is associated with working memory.[52] The high
gamma activity is spatially and temporally different depending on the movement type.[40,53,54] Moreover, in the motor imagery task, the spatial distribution of neural activity at high frequencies (76–100 Hz) and low frequencies (8–32 Hz) was similar to the activation pattern during actual motor movement
(Figure 2c).[41] For the language task, Towle et al. showed that
the gamma activity varies spatially depending on listening and
speaking tasks. During speaking, compared to listening, activity in the frontal lobe increased while activity in the temporal
lobe decreased.[55] Another study on cortical connectivity revealed
that the role of Broca’s area in language appears to formulate articulatory codes for speech pronunciation.[56] In addition, studies have suggested that neural activity is modulated by various
cognitive activities such as auditory and visual imagery tasks
(Figure 2d).[42,57]
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Figure 2. Examples of human ECoG studies in medicine and neuroscience fields. a) Representative ictal discharges across multiple electrodes (top) and
dynamic voltage topographic maps on the cortical surface at the time points I to IV (bottom).[39] b) Time-frequency maps for various vibrotactile stimulus
conditions in the primary somatosensory cortex (S1) and secondary somatosensory cortex (S2).[49] c) Power spectral density of cortical potentials for
motor execution (left, red line), motor imagery (right, red line), and resting state (both, blue line) in the primary motor cortex.[41] d) Change in high
gamma power during music listening relative to the baseline in the posterior superior temporal gyrus.[42] Reproduced with permission from publishers.
2.2.3. Brain–Computer/Machine Interface (BCI/BMI)
Based on the neuroscientific findings discussed in the previous
section, studies utilizing ECoG for BCI (or BMI) applications
have recently emerged. Previous BCI studies have extensively
utilized information on spike firing obtained through intracortical recording, which is a more invasive method than ECoG.[58–60]
However, recent efforts have been made to use ECoG to decode
brain activities for BCI purposes, demonstrating the feasibility
of using ECoG in decoding movement or communication intentions. For instance, for motor BCI, Kang et al. suggested a
method for expecting movement intention using the temporal
dynamics of brain connectivity in the dorsolateral prefrontal and
primary motor cortices.[61] Talakoub et al. predicted the trajectory of real arm movements with high performance.[62] Jang et al.
demonstrated the successful prediction of imagined hand movement trajectories and suggested a more appropriate paradigm
for higher decoding accuracy.[63] In addition, Lorach et al. iden-
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tified hip flexion using ECoG from the sensorimotor cortex to
restore communication between the brain and spinal cord from
patients with chronic tetraplegia.[64] Another study demonstrated
the feasibility of decoding hand gestures with high accuracy from
high-density ECoG signals in the sensorimotor cortex, highlighting the potential for developing a high-performance BCI using
high-density ECoG.[54]
For the auditory BCI system, Martin et al. demonstrated the
feasibility of a music BCI system by suggesting an encoding
model to predict the high gamma activity (70–150 Hz) from the
auditory spectrogram during the music imagery task.[65] Anumanchipalli et al. succeeded in designing a speech decoder using cortical activity related to articulatory movement.[66] A recent
study showed that direct cortical signals from high-density ECoG
electrodes in the sensorimotor cortex could be used to decode
words and sentences during overt and silent speech in real-time
with high accuracy for a paralyzed participant to control communication BCI.[67,68]
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Table 1. Commercial ECoG devices that have been used in human subjects.
Reference
Pacia et al., 2022[ 76]
Inoue et al., 2021[ 77]
Sommer et al., 2018[ 78]
Chiba et al., 2022[ 79]
Kim et al., 2016[ 80]
strip/grid
Type
strip/grid
strip/grid
strip/grid
strip/grid
Number of Subjects
21
9
13
25
10
Implantation period (day)
2-13
10
3-7
3-14
6-11
Diameter (mm)
4
5
6
5
5
Exposure size (mm)
2.3
3
5
3
3
Center-to-center spacing
(mm)
10
10
10
10
10
Material
Platinum or stainless
steel
Platinum
Platinum
Platinum
–
Substrate
Medical silicone
Medical silicone
Medical silicone
Medical silicone
Medical silicone
2.3. Commercial ECoG Electrodes
Macro-sized ECoG electrode arrays have been commercialized
and have been widely used in hospitals since the first Palm Desert
Conference on the Treatment of Epilepsy in 1986.[69] Commercial
ECoG arrays are generally used for the invasive monitoring of patients with epilepsy in clinical practice, mainly within 1–2 weeks,
and are implanted for a maximum of 30 days (Table 1). In general,
ECoG signals allow for the recording of extracellular field potentials and the study of the connectivity of brain regions, allowing
us to understand various pathological responses and cognitive
processes. In addition, ECoG electrodes have a high spatial resolution at the millimeter scale. In addition, since dozens to 200
electrodes are generally implanted for invasive monitoring, a relatively broad area of the cortex is covered. Moreover, electrical
stimulation can be performed directly on a specific cortical area
using ECoG electrodes.[70]
Commercial ECoG arrays are mostly composed of a medical
silicone substrate to minimize foreign body responses and have
conductive parts such as electrodes, traces, and pads made of
stainless steel or platinum to minimize chemical reactions with
biological tissues. Platinum electrodes are preferred for their superior resistance to electrolysis and stable electrochemical properties, making them more suitable for long-term applications and
consistent signal quality.[71] Conversely, stainless steel electrodes
are more cost-effective, though they are less resistant to corrosion and may subsequently exhibit less stable electrical properties for chronic applications.[72] An ECoG array consists of a strip
or a grid of electrodes (Figure 3). In general, a strip has 2–12 electrodes arranged in a line, and a grid has 8–64 electrodes arranged
Figure 3. Examples of commercial ECoG electrodes used in the clinical field according to manufacturers. a) g.tec Medical Engineering GmbH, b) PMT
Corporation, and c) Ad-Tech Medical Instrument Corporation, respectively. An electrode array with multiple rows and columns is referred to as a grid-type
electrode, while an electrode array with a single row is referred to as a stripe-type electrode.
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in a grid. The size of the metal electrodes ranges from 3 to 6 mm,
and the size of the exposure is slightly smaller. It is estimated
that ≈500 000 cells exist near an electrode or in the local area underneath it.[73] Generally, the electrode spacing ranges from 5 to
15 mm, and the thickness of the ECoG electrode arrays is usually
0.7 mm. The 8 × 8 grid covers a relatively wide cortical area of 80
× 80 mm2 . The size and shape, such as the L-shape or T-shape,
and the number of electrodes can be customized if needed. An
electrode with a contact size ˂2 mm and an inter-electrode distance ˂5 mm is regarded as a high-density electrode. Such highdensity ECoG arrays can even detect high-frequency ripples of
hundreds of hertz in addition to normal ECoG signals during
seizures.[74]
2.4. Limitations of Commercial ECoG Electrodes
2.4.1. Mechanical Mismatch
The substrate of commercial ECoG arrays is typically medical
silicone, which is a soft elastomer. Even though the substrate
material is soft, millimeter-thick substrates possess high bending stiffnesses. The electrodes are made of rigid stainless steel
or platinum plates. Thus, the mechanical properties of commercial ECoG arrays are much stiffer than soft brain tissues. The
stiffness of ECoG arrays encounters inherent problems during
the implantation and recording of electrophysiological signals.
Stiff ECoG arrays do not have good contact with curved and
soft brain tissues. Consequently, it hinders high-quality ECoG
recordings. In addition, risks including symptomatic extra-axial
fluid collection may be reduced by reducing the stiffness of
the electrode arrays.[75] However, the manufacturing method of
commercial ECoG electrodes has limited the fabrication of soft
ECoG arrays owing to the procedures used and the low yield in
production.
2.4.2. Low Spatial Resolution
The spatial resolution of electrophysiological signals is determined by the electrode size and density. The millimeter-sized
electrodes and electrode arrangement of commercial ECoG arrays limit the spatial resolution as they can record the population
activities of neurons close to the electrodes. To achieve a higher
spatial resolution, electrodes need to be fabricated on a submillimeter scale with high-density to capture the activity of a smaller
population of neurons.
2.4.3. Limited Functionality
The additional functions of ECoG arrays can expand the potential for target-specific drug delivery, neural circuit mapping, and
minimal surgical risk. Commercial ECoG arrays focus on measuring electrophysiological signals from the human brain. Therefore, commercial ECoG arrays are limited for use in small animals, such as mice, rats, or monkeys, where brain mechanisms
can be explored more actively. In addition, investigating neural
circuits by measuring only electrophysiological signals is challenging. For an in-depth understanding of neural circuit mechanisms, it is desirable to integrate ECoG arrays with systemic
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modalities such as microfluidic or optical components. Furthermore, commercial ECoG arrays implanted after cranial incision
require a large exposure window of 28 ± 15 cm2 on average,[81]
which can cause adverse effects such as clinical infection, subdural hemorrhage,[82–86] cerebral edema, or aseptic meningitis.
Minimally invasive surgical and systemic strategies are expected
to reduce surgical side effects in patients.[86,87]
3. Development of Micro-ECoG Electrodes
Since its introduction by Jasper and Penfield in 1949, ECoG
electrodes have been commonly used to directly measure neural activity on the cortical surface in the clinical and research
fields.[88,89] Since the 2000s, microelectromechanical system technology has been actively adopted to fabricate ECoG electrodes,
and subsequently, an understanding of brain activity at high spatial resolution has become possible (Figure 4a).[90] In particular, polymer materials have been used as substrate materials in
the fabrication of ECoG electrodes (Figure 4b). Some soft ECoG
electrodes can establish conformal contact between the electrode
and brain tissue.[91,92] Recently, ECoG electrodes have also been
integrated with various additional functionalities for the study
of the complex mechanisms of the brain and the development
of neurological disease treatment. For example, optical imaging and optogenetic stimulation have been integrated into ECoG
electrodes for research purposes.[93,94] A drug-delivery function
was added for direct and focal injection into the central nervous
system.[95] Bioabsorbable materials and minimally invasive electrode array structures have been developed to reduce surgical
risks.[91,96]
3.1. Improvement of Mechanical Properties
The substrates of ECoG electrodes are recommended to be made
of soft polymeric materials to reduce the mechanical mismatch
between the electrodes and the brain. Thus, biocompatible polymeric materials, such as plastics, elastic materials, and hydrogels,
are used as substrates (Figure 5). The Young’s moduli of plastic
materials,[2,97–101] elastic materials,[102–104] and hydrogels[92] are
typically ≈10 GPa, ≈10 MPa, and ≈1 MPa, respectively. Therefore, to fabricate ECoG electrodes with mechanical properties
similar to those of brain tissues, the substrate materials should
be considered in terms of Young’s moduli and their suitability for microfabrication (Table 2). Additionally, to represent the
flexibility of thin film ECoG devices, the bending stiffness (e.g.,
flexural rigidity, K) is one of the important parameters to be
considered.[105]
K =
F
48EI
= 3
d
L
(1)
The bending stiffness of the beam-structured electrode is determined as the bending force that is needed to achieve conformal contact with the curved cortex surface. Here, F is the applied
force, d is the deflection, E is Young’s moduli of substrate materials, I is the moment of inertia, and L is the length of the
electrode. The cross-sectional shape of the ECoG electrode is
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Figure 4. Trends of publications regarding ECoG electrodes from 1970 to 2021. a) The blue column indicates publication trends using the term “ECoG
electrode”, among which studies searched using the terms “micro” and “ECoG electrode” are shown in magenta, and b) studies searched using the
terms “flexible” or “soft” together with “ECoG electrode” are shown in green. All data are based on Scopus.
normally rectangular shape. The formation of the moment of inertia is expressed as
K =
48Ewh3
L3
If the ECoG electrode is assumed to be a wide rectangular sheet
(h ≪ L), the bending stiffness of the beam-structured electrode
can be expressed as
(2)
K =
where w and h are the width and thickness of the electrode substrate, respectively.
Eh3
12 (1 − 𝜈 2 )
(3)
where 𝜈 is Poisson’s ratio.
Figure 5. Representative substrate materials for ECoG electrodes. a) Young’s modulus of the brain tissue to materials used in ECoG electrodes. b)
Optical images of substrate materials such as hydrogel (b-1), polydimethylsiloxane (PDMS) (b-2), medical-grade silicone used for commercial ECoG
electrodes (b-3), and parylene C (b-4) when placed on the model of the human brain; (top) aerial- and (bottom) side views. The red dotted lines indicate
the outline of the curved brain surface. The green dotted lines indicate the outline of each substrate material placed on the curved brain surface. The
gap between the brain surface and the substrate is marked by black arrows.
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Table 2. Properties of the common materials used to build ECoG electrode arrays.
Electrode
Polymeric substrate
Metal
Plastic material
Elastic material
Hydrogel
PDMS[ 107,137–140]
PVA[ 141–145]
Au
PI[ 107,131–133]
Parylene C[ 107,134–136]
Young’s modulus (MPa)
60 000
2500
3200
1.35
0.066
Melting temperature (°C)
1064
–
290
−46.3
220–230
Glass transition temperature (°C)
–
276–302
80–100
−123
46
Thermal coefficient of Expansion
(ppm/K)
14.2
12
35
300
40
Thermal Conductivity (W/cm K)
3.18
0.29
8.2
15–25
0.34
Water vapor permeability
(g/m2 /day)
–
0.867–8.24
3.26–3.88
432
28 600
Biocompatability
–
–
USP calss VI
USP calss VI
FDA approved
Physical and biological properties
PI, Polyimide; PDMS, Polydimethylsiloxane; PVA, polyvinyl alcohol; USP, U.S. Pharmacopeia; FDA, Food and Drug Administration. USP calss VI indicates there are no harmful
reactions or long-term bodily effects caused by chemicals. FDA approval means that the hydrogel is permitted for clinical use.
The device structure of thin film enhances flexibility and enables conformal contact with the curved surface such as the brain
cortex.[106] Therefore, conductive components such as electrodes
and conductive lines are required to form flexible structures using microfabrication techniques. Plastic materials have higher
Young’s moduli than elastic materials (elastomers) and hydrogels. These materials have been widely employed to fabricate
ECoG electrodes owing to their convenience and compatibility
with microfabrication techniques. Elastomers and hydrogels are
excellent substrate materials in terms of mechanical properties
similar to those of the brain tissue. However, these materials
exhibit low suitability for microfabrication techniques. A lot of
ECoG electrodes are fabricated based on diverse substrate materials and have been developed to achieve stable chronic recording
from the brain cortex (Table 3).
3.1.1. Plastic Materials
Plastic materials possess relatively higher Young’s moduli among
the polymeric substrates of the ECoG arrays. To reduce the
Table 3. Summary of the developed ECoG electrodes for in-vivo recording studies.
Substrate
PI
Parylene
PDMS
Hydrogel
Thickness
Subject
Implantation area
Duration
Main outcome
Ref.
25 μm
Cat
Visual cortex
Acute
High density
[111]
10 μm
Macaque monkey
Hemisphere
18 weeks
Free movement of animal
[146]
8 μm
Cat
Visual cortex
16 days
Chronic optogenetics
[119]
25 μm
Mouse
Somatosensory cortex
12 weeks
Recording from thinned skull
[108]
7 μm
Rat
Visual and motor cortex
2 weeks
Minimal skull opening
[81]
7 μm
Rat, human
Barrel cortex
Acute
High density without cross-talk
[118,147]
20 μm
Macaque monkey
Somatosensory cortex
Acute
High density
[100,117]
4 μm
Rat, human
Neocortex
10 days
High density
[2, 116]
13 μm
Rat
Cortex
Acute
Flexibility
[148]
80 μm
Macaque monkey
Somatosensory cortex
12 weeks
Long-term recording
[105]
200 μm
Rat, cat, zebrafish
Sensorimotor cortex
6-8 weeks
Stretchable and rapid fabrication
[149]
20 μm
Rat
Somatosensory cortex, parietal
association cortex, and visual
cortex
Acute
Multi-cortical brain
[124]
75 μm
Rat
Somatosensory cortex
13 weeks
Stretchability
[104]
200 μm
Rat
Auditory cortex
Acute
Cost-effective and simple
fabrication
[150]
30 μm
Rat
Barrel cortex
33 days
Bioresorbablity
[151]
40 μm
Rat
Cerebral cortex
5 days
Bioresorbablity
[96]
20 μm
Rat
Cerebral cortex
20 weeks
Chronic optogenetics
[129]
1000 μm
Porcine
Cortex
Acute
Highly soft and oxygen nutrition
[128]
5 μm
Mouse
Somatosensory and motor cortex
Acute
Transparent
[152]
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mechanical mismatch between biological tissues and plastic substrates, the substrate is fabricated with a thickness of down to tens
of micrometers using chemical vapor deposition (CVD) or spincoating procedures.
Polyimide (PI) has a coefficient of thermal expansion similar to
that of silicon. Thus, it is compatible with most microfabrication
techniques that are conducted on silicon wafers. The glass transition temperature is higher than the processing temperature during metal electrode fabrication. Owing to these characteristics, PI
is widely utilized in the fabrication of microelectronics. Moreover,
advantages such as superior corrosion resistance and thermoxidative stability are the reasons for PI being preferred as an insulation and passivation layer.[107] Due to the compatibility with microfabrication techniques, a variety of PI-based ECoG electrodes
have been developed at the micrometer scale.[81,101,108–112]
Parylene C, a plastic material named polyparaxylene C, exhibits desirable characteristics for implantation in terms of
chemical inertness, ion/moisture barrier function, optical transparency, and medical-grade biocompatibility (U.S. Pharmacopeia
(USP) class VI approved by Food and Drug Administration
(FDA)). Parylene C can be conformally and uniformly deposited
as a substrate or as an insulation material for implantable devices. However, the CVD-based deposition limits the thickness
of deposited parylene C. Parylene C also exhibits high compatibility with microfabrication techniques except for the procedures
that are conducted at high temperatures due to its glass transition temperature of 80–100 °C. Thanks to parylene C exhibiting high compatibility with microfabrication techniques, many
ECoG electrodes based on parylene C substrates have been fabricated and applied to animal or human studies.[2,99,102,113–118]
3.1.2. Elastic Materials
Elastic materials are soft, stretchable polymers that do not undergo plastic deformation. Polydimethylsiloxane (PDMS) is a
popular elastic material with the advantages of optical transparency, biocompatibility, and resistance to biodegradation. In
particular, PDMS is a material approved to be USP class VI and
has mechanical properties (Young’s modulus close to 1 MPa)
similar to those of biological tissues (Young’s modulus of the
brain ≈100 kPa).[119] However, this softness compromises the
compatibility with microfabrication techniques owing to the
large difference in the thermal expansion coefficient between
the metals and PDMS. Recently, a fabrication method for stable
metal patterns on a soft PDMS substrate was established using
an intermediate parylene C layer,[120] which improved the compatibility of PDMS with microfabrication techniques.[121,122] Various efforts have been made to utilize PDMS substrates for ECoG
electrodes.[95,102,104,123–125]
3.1.3. Hydrogel
Hydrogels have excellent mechanical properties similar to those
of biological tissues and have attracted broad attention as
biomaterials for applications such as electronic skins[126] or
bioelectronics.[127] Hydrogels possess additional advantages such
as transparency, biofluid permeability, and adhesion to biologi-
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cal tissues. Electrical functions can be integrated by mixing hydrogels with conductive nanowires or nanocomposites. However,
hydrogel-based substrates are incompatible with general microfabrication techniques because of the changes in their mechanical properties, such as volume expansion and decreased Young’s
moduli, upon water absorption. Thus, alternative methods are required to integrate conductive patterns onto hydrogel substrates.
Recently, polyvinyl alcohol (PVA) hydrogel substrate-based ECoG
electrodes have been developed to enhance the conformal contact
with the brain surface and simultaneously record ECoG signals
during fMRI.[128]
Recent studies report that conductive patterns based on hydrogels were fabricated by mixing PVA with Ag/Au nano core-shell
nanowires[129] and polystyrene sulfonate (PEDOT:PSS)[130] to enhance the long-term stability of ECoG electrodes in aqueous environments. They recorded ECoG signals from the cerebral cortices
of marmosets and mouse brain slices.
3.2. Improvement of Spatial Resolution
Efforts to communicate with the brain have encouraged the
use of high spatiotemporal resolution to record neural activity
with spike-level accuracy in both BCI/BMI research and clinical seizure tracking (Figure 6). Physiological signals from microECoG (μECoG) arrays can improve our understanding of physiological and pathological mechanisms with higher resolutions.[70]
A micro-ECoG array with a sampling rate of 20–30 kHz was
shown to record brain activity in a wide frequency band, up
to single-unit activities.[2,116] Accordingly, to decode brain functions such as perception, sensation, and movement, microelectrodes with a small size under 100 × 100 μm2 and high density are preferred to detect the activity of individual neurons.
As an example, Buzsaki’s group fabricated a micro-ECoG array,
named “NeuroGrid,” with a 10 μm-diameter and 23 μm-spacing
of electrodes.[116] NeuroGrid demonstrated the acquisition of action potential waveforms from superficial cortical neurons with
high spatiotemporal resolution.[2]
Moreover, if micro-ECoG electrodes are used in clinical fields,
precise seizure tracking can be achieved by accurately detecting ictal onset areas and identifying epileptic pathways in the
core and penumbra areas. In addition, ECoG signals with
high spatiotemporal resolution can improve the controllability
of BCI/BMI. Dayeh’s group developed 2048 ECoG electrodes
named “PtNRGrids” to spatially confirm epileptic discharge
spreads from a patient.[118] Also, they recorded complex spatiotemporal activities from the somatosensory and motor cortices
of awake humans, associated with the vibration of fingers and
grasping task, respectively.
Micro-ECoG arrays have been fabricated using microfabrication techniques to achieve high spatial resolution (Figure 7a).
Microfabrication techniques enable not only the deposition and
patterning of metal thin films down to the submicrometer scale
but also the coating and etching of various organic/inorganic materials. Therefore, it provides choices for the material selection
in the fabrication of ECoG electrodes. Recently, laser machining,
inkjet printing, and 3D printing have been used to rapidly and
conveniently develop ECoG electrodes (Figure 7b–d).[102,149,156,157]
Although conventional microfabrication techniques can
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Figure 6. Comparison between macro- and micro-ECoG array (ECoG versus micro-ECoG). The schematic shows the characteristics of ECoG and microECoG. The ictal onset zone is the specific area where a seizure starts. The core surrounding the ictal onset zone can be defined by detecting brain
activities using a micro-ECoG array with higher spatial resolution than an ECoG array. The penumbra, defined by using an ECoG array, refers to the
surrounding area including the central core of a seizure zone. Intense and synchronized brain activities are observed in the seizure core, while low-level
and desynchronized firings are observed in the penumbra.[153–155]
implement metal electrodes in tens of micrometers, it has
disadvantages such as complex manufacturing procedures, and
ineffectiveness in time and cost (Figure 7e). Other methods of
laser machining, inkjet printing, and 3D printing have advantages such as simple fabrication procedures, but they imply
relatively low resolution of electrode patterns and limited materials to be employed. Therefore, it is necessary to understand the
available electrode materials based on the fabrication method.
Consequently, a suitable fabrication strategy for ECoG electrodes
is required for specific clinical applications and neuroscientific
research.
3.2.1. Microfabrication
Microfabrication techniques for ECoG electrodes such as photolithography, sputtering, evaporation, and dry/wet etching
have enabled the deposition of various metal thin films of
Au,[102,109–111,113,150,160,173] Pt,[108,114,146,174] and Ir,[98,161,170] with the
size of electrode patterns ranging from 10 to 1000 μm. In
particular, Au and Pt have generally been selected as electrode materials for ECoG electrodes because of their high conductivity, biocompatibility, and chemical inertness. In addition,
indium tin oxide (ITO), graphene, and conductive nanowirebased ECoG electrodes have been developed owing to their
transparent characteristics to integrate bioimaging and optogenetic functions.[112,115,129,162,163,172,175] Moreover, the electrical performance of micro-sized electrodes can be improved
by coating specific materials onto a metal thin film. For example, IrOx,[98,170] poly(3,4-ethylenedioxythiophene) polystyrene
sulfonate (PEDOT:PSS),[2,99,116,147,152] carbon nanotubes,[124] Pt
black,[81,117,164] and metal-PDMS composite[95,104,123,165] have been
employed to enhance electrode surfaces (Figure 7f). Consequently, the microfabrication of ECoG electrodes is versatile
in terms of material selection for achieving high-quality signal
recording. This enables electrode patterning on a scale of sev-
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eral micrometers with high stability and yield in manufacturing
processes.[2,81,95,98,100,102,108,112–115,117,123,124,146,147,150,160–163,165,166,172]
3.2.2. Laser Machining
The fabrication procedures of ECoG electrodes through laser machining are similar to those of microfabrication, except for the
patterning of the metal film and/or array outline. This method is
relatively convenient and fast since laser cutting is used instead
of a photolithography-dry/wet etching-cleaning cycle. A Pt or PtIr
foil is laminated onto the substrate and patterned to create electrodes with a resolution of several hundreds of micrometers using an ND:YDG laser.[103,176] However, laser machining has several challenges, such as low cutting resolution caused by damage
from laser ablation and limited employable materials. Thus, this
method limits the fabrication of micro-ECoG electrodes to resolutions of several micrometers.
3.2.3. Inkjet Printing
Inkjet printing has been developed using drop-on-demand patterning technology to directly disperse conductive materials on a
scale of tens of micrometers. This maskless patterning method
has the advantages of low cost and short throughput in manufacturing. However, solvent- or water-soluble materials, such as Ag
ink and PEDOT:PSS, which can be ejected through an inkjet nozzle, are necessary for the fabrication of microelectrodes.[148,158]
Recently, inkjet printing-based electrodes with a diameter of
30 μm were fabricated, so that they could have a resolution sufficient for single-cell recording.[157]
3.2.4. 3D Printing
3D printing has contributed to the convenient, simple, and rapid
fabrication of ECoG arrays, by printing conductive materials of
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Figure 7. Fabrication strategies and properties of ECoG electrodes. a-d) Images of ECoG arrays fabricated using various manufacturing techniques,
along with representative electrode materials used in each fabrication method: a) microfabrication,[102] b) laser machining,[156] c) inkjet printing,[157]
and d) 3D printing.[149] e) Resolution versus throughput time according to fabrication techniques.[2,81,95,98,100,102,103,108,112–115,117,123,124,146–150,157–168]
Filled symbols represent the highest resolution that can be achieved in each fabrication strategy. f) Electrode impedance at 1 kHz according to electrode
materials. Figures a to d are reproduced with permission from publishers.[2,95,98,100,102,104,108,112–115,117,123,124,146,147,149,150,160–162,164–166,169–172]
PEDOT:PSS with patterning resolutions of 100–200 μm.[149,177]
Recently, inkjet printing of Ag ink was converged with 3D printing of guide structure.[149] The substrates and conductive electrodes have been fully printed in a one-step process such that the
printing method enhances the advantages of fabricating ECoG
arrays rapidly and conveniently.
functions and a promising approach to treating neurological diseases. Thus, micro-ECoG arrays have been integrated with various functions, including drug delivery, optical stimulation, fluorescence imaging, minimally invasive implantation, and bioabsorbable materials (Figure 8). ECoG arrays with additional functions are being developed to be more biocompatible in biological
environments, deeply understand brain mechanisms, and meet
clinical requirements.
3.3. Integration of Additional Functions
Previously, the role of micro-ECoG arrays was limited to measuring electrophysiological signals or modulating neural activities. However, recent results from neuroscience research and
clinical applications have shown that multifaceted studies are
required for an in-depth understanding of higher-order brain
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3.3.1. Drug Delivery
ECoG electrodes have been integrated with drug delivery functions for precise and target-specific drug infusion. In a study by
Minev et al., a microfluidic channel passively delivered drugs via
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Figure 8. Functionalities added to micro-ECoG arrays. ECoG arrays are integrated with a) microfluidic channels for drug delivery,[95] b) light emitting component for optogenetic modulation,[129] c) transparent substrate/electrode for optical monitoring,[178] d) expandable structure for minimal
invasiveness,[179] e) biodegradable materials for bioabsorbability.[130] Reproduced with permission from publishers.
mechanical movement to eliminate the side effects of serotonergic agents on the autonomic system.[95] A flexible drug delivery
microdevice (f-DDM) released an antiepileptic drug (AED) with
a controllable amount and multiple drugs for the precise control of seizures.[180] Especially, an electrophoretic ion pump was
included to resolve the issues of clogging and reflux, thereby enhancing drug delivery performance.[181] In another study, bioresorbable silk fibroin was used to deliver phenobarbital, an AED.
As a result, the suppression of epileptic seizures was observed
with the dissolution of silk fibroin.[110] These studies were conducted to enhance the performance of drug delivery in precision
medicine.
3.3.2. Optogenetic Modulation
Optogenetic modulation is a prospective method for investigating the functional connectivity of neural circuits with a high spatiotemporal resolution. Using this method, a circuit connection
with specific cell-type activity can be identified. The stimulation
artifact, which is usually seen in electrical stimulation, is eliminated; therefore, the optogenetic method has a definite advantage
in neuromodulation studies.[182] For stimulation and recording
in the same brain area, the electrodes and substrate of a microECoG device need to be composed of transparent materials. Accordingly, transparent ECoG arrays fabricated using various electrode materials such as ITO, graphene, carbon nanotubes, metal
nanowires, or PEDOT:PSS have been developed.[115,172,183,184] An
ECoG array including a transparent ITO electrode was integrated
with the micro-LED array for optical modulation and simultaneous recording of induced signals in the visual cortex.[185] Embed-
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ded micro-LED arrays can simultaneously achieve optical stimulation in multiple areas of the cortex.
3.3.3. Optical Monitoring
Integrating an optical window and transparent ECoG array can
facilitate simultaneous optical imaging and electrophysiological
signal recording. Simultaneous detection using both modalities
allows multifaceted studies of neural circuits. Au nanospheres
were used as the transparent electrode material and exhibited
an average transmittance of over 65% at 300–1100 nm light
wavelength.[186] In this study, the cell morphology was successfully captured through an Au nanosphere electrode by phasecontrast imaging. In addition, graphene-based ECoG electrodes
with a transmittance of over 90% were fabricated to acquire
spatiotemporal neural fluorescence responses.[175] A multilayerassembled ECoG array with PEDOT:PSS, ITO, and Ag, resulting
in a transmittance of 89% at 550 nm, was developed to conduct
electrical and optical multimodal recording.[187]
3.3.4. Minimal Invasiveness
The implantation of ECoG electrodes generally requires a larger
opening of the skull than the lesion site, accompanying patient
risks, such as a high probability of infection and brain swelling.
Therefore, several solutions have been devised to reduce the size
of the trephination hole but keep the contact area of ECoG arrays
as needed. An expandable ECoG array, composed of an elastic
PDMS substrate, was suggested to implant in the gap between
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the skull and cortex through a smaller trephination hole than the
cortical area to cover, by expanding the array inside the gap.[179]
Another ECoG electrode integrated with a neodymium magnet
was developed, and the placement of the device at a desired location inside the skull was demonstrated.[81] After positioning the
device, the magnet utilized as the carrier was released from the
device and extracted through the hole in the skull. Recently, a
soft robotic electrode, which can be expanded and cover a large
area of cortical surface by hydraulic injection, demonstrated minimally invasive cortical surgery for applications of neurological
disorders.[188]
Acknowledgements
3.3.5. Bioabsorbability
Keywords
Bioresorbable ECoG electrodes can reduce the postoperative risk
of surgical extraction of the device after a certain period of
time. Polycaprolactone (PCL)- and PCL/ poly(lactic-co-glycolic
acid) (PLGA)-based electrode substrates were fully fabricated by
inkjet printing, and bioabsorbability was confirmed in accelerated aging environments for up to 70 h.[157] This result is projected for a degradable period of two years in vivo, confirming
the slow degradation of the implanted devices. In addition, the
PLLA/PCL substrate and molybden (Mo)-based ECoG array were
fabricated through microfabrication techniques and dissolved in
37 °C phosphate buffered saline solution for 30 days. The ECoG
array detected the onset of epileptic seizures.[96] Moreover, the
biodegradable Si nanomembranes (Si NMs)-based ECoG arrays
achieved chronic recordings of seizure-like spiking activity for 30
days.[151]
electrocorticogram (ECoG), ECoG electrodes, micro-ECoG
4. Challenges and Prospects
As discussed in this review, medical and neuroscientific research
on ECoG has expanded remarkably since 2010. The majority
of EcoG usage is still observed in clinical monitoring and diagnosis; however, its use in neuroscientific discovery and technological development, such as in BCI/BMI, is rapidly increasing. Along with these trends, micro-ECoG devices and flexible
and/or soft ECoG devices have been actively developed with improved biocompatibility, higher spatial resolution, and additional
functionalities owing to the advancements in materials and fabrication technologies. These new developments are currently
used in animal models in combination with neuroscientific studies. This needs to be translated into clinics to advance clinical
ECoG toward more personalized and customized treatments and
next-generation precision medicine, such as electroceuticals. The
main aim of micro-ECoG devices is the achievement of advanced
healthcare for patients with neurological diseases and the advancement of neural interfaces with reliable long-term performance. Moreover, the devices aim to provide minimal invasiveness for severely paralyzed patients with impaired physical functions due to spinal cord injury, amyotrophic lateral sclerosis, or
locked-in syndrome, as well as for patients with cognitive impairment. Finally, in parallel with the efforts to develop new technologies, efforts to obtain approval from the regulatory entity in each
country should be present simultaneously to realize the translation of promising new technologies.
Adv. Mater. Technol. 2024, 9, 2301692
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H.M., J.K., and J.E. contributed equally to this work. This work was
supported by grants of the National Research Foundation (grant
no.: NRF-2018M3C7A1022309, NRF-2018M3C7A1022317 and NRF2021R1C1C2011489) and Korea Brain Research Institute (grant no.:
24-BR-02-02, 24-BR-03-04 and 24-BR-04-03), funded by the Ministry of
Science and ICT of Korea.
Conflict of Interest
The authors declare no conflict of interest.
Received: October 8, 2023
Revised: February 14, 2024
Published online: March 21, 2024
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Hyunmin Moon received his BS degree in biomedical engineering from Gachon University, Korea, in
2012. He received MS and PhD degrees in robotics engineering from Daegu Gyeongbuk Institute of
Science and Technology (DGIST), Korea, in 2014 and 2021, respectively. From 2021 to 2023, he was a
postdoctoral researcher with the Department of Robotics and Mechatronics Engineering at DGIST.
Since 2023, he has been a postdoctoral researcher with the Department of Mechanical Engineering at
Massachusetts Institute of Technology (MIT), USA. His research interests include neural interfaces
for brain and peripheral nerve applications as well as soft bioelectronics employing robust, conductive, and bioadhesive hydrogels.
Jii Kwon received her BS degree in nano-physics from Gachon University, Korea, in 2018. She is
presently an integrated MS and PhD student in the Department of Brain Cognitive and Science at
Seoul National University, Korea. Her research is primarily focused on brain signal decoding and stimulation for auditory brain-computer interfaces and epilepsy.
Jonghee Eun received her BS degree in physics from Chungbuk National University, Korea, in 2016.
She received her PhD degree in physics from Ulsan National Institute of Science and Technology
(UNIST), Korea, in 2021. After receiving her PhD, she worked as a postdoctoral researcher at UNIST
until 2022. In 2022, she joined as a postdoctoral researcher with the Emotion, Cognition & Behavior
Research Group at Korea Brain Research Institute (KBRI), Korea. Her current research interests include the development of multifunctional neural interface platforms for neuroscience and ultrasound
applications for neurodegenerative diseases.
Chun Kee Chung received his MD and PhD degrees from Seoul National University (SNU), Korea.
He was affiliated with the Department of Neurosurgery, SNU College of Medicine and Hospital as a
professor from 1993 to 2023. After retirement, he is currently a senior researcher at the Neuroscience
Research Institute at SNU College of Medicine. His research interests include brain–machine interfaces, epilepsy, memory, language, sensorimotor interaction, and consciousness.
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June Sic Kim received his BS degree in electrical engineering, and MS and PhD degrees in biomedical
engineering from Hanyang University, Korea, in 1996, 1998, and 2002, respectively. He worked with
Montreal Neurological Institute, McGill University, Canada, from 2002 to 2004, as a postdoctoral
researcher. He worked as an assistant professor with the Department of Neurosurgery, Seoul National
University Hospital (SNUH), Korea, from 2004 to 2014. He worked as a research professor with the
Institute of Basic Sciences at SNU from 2014 to 2023. Since 2023, He has been working at Konkuk
University Medical Center as a research professor. He is interested in brain-computer interfaces based
on neuroimage and electrophysiology.
Namsun Chou received his BS degree in mechanical engineering from Konkuk University, Korea, in
2009. He received MS and PhD degrees in mechatronics from Gwangju Institute of Science and Technology (GIST), Korea, in 2011 and 2016, respectively. He worked with the Department of Robotics
Engineering at DGIST, from 2016 to 2017, and the Brain Science Institute at Korea Institute of Science
and Technology (KIST) from 2017 to 2021, as a postdoctoral researcher. Since 2021, he has been a
senior researcher with the Emotion, Cognition & Behavior Research Group at Korea Brain Research
Institute (KBRI). He is interested in multifunctional neural interfaces for neuroscience studies.
Sohee Kim received her BS and MS degrees in mechanical engineering from KAIST, Korea, in 1998
and 2000, respectively, and PhD degree in mechatronics from University of Saarland, Germany, in
2005. From 2006 to 2009, she was a postdoctoral researcher in electrical and computer engineering at
University of Utah, USA. From 2009 to 2015, she was a professor at GIST, Korea. Since 2015, she has
been a professor with the Department of Robotics and Mechatronics Engineering at DGIST, Korea.
Her research interests include neural interfaces for brain and peripheral nerve applications as well as
polymer-based soft MEMS technologies.
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