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EEGNet: A Compact Convolutional Network for EEG-based Brain-Computer Interfaces


Abstract and Figures

Objective: Brain computer interfaces (BCI) enable direct communication with a computer, using neural activity as the control signal. This neural signal is generally chosen from a variety of well-studied electroencephalogram (EEG) signals. For a given BCI paradigm, feature extractors and classifiers are tailored to the distinct characteristics of its expected EEG control signal, limiting its application to that specific signal. Convolutional Neural Networks (CNNs), which have been used in computer vision and speech recognition to perform automatic feature extraction and classification, have successfully been applied to EEG-based BCIs; however, they have mainly been applied to single BCI paradigms and thus it remains unclear how these architectures generalize to other paradigms. Here, we ask if we can design a single CNN architecture to accurately classify EEG signals from different BCI paradigms, while simultaneously being as compact as possible. Approach: In this work we introduce EEGNet, a compact convolutional neural network for EEG-based BCIs. We introduce the use of depthwise and separable convolutions to construct an EEG-specific model which encapsulates well-known EEG feature extraction concepts for BCI. We compare EEGNet, both for within-subject and cross-subject classification, to current state-of-the-art approaches across four BCI paradigms: P300 visual-evoked potentials, error-related negativity responses (ERN), movement-related cortical potentials (MRCP), and sensory motor rhythms (SMR). Results: We show that EEGNet generalizes across paradigms better than, and achieves comparably high performance to, the reference algorithms when only limited training data is available across all tested paradigms. In addition, we demonstrate three different approaches to visualize the contents of a trained EEGNet model to enable interpretation of the learned features. Significance: Our results suggest that EEGNet is robust enough to learn a wide variety of interpretable features over a range of BCI tasks.Our models can be found at:
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EEGNet: A Compact Convolutional Neural Network
for EEG-based Brain-Computer Interfaces
Vernon J. Lawhern1,*, Amelia J. Solon1,2, Nicholas R. Waytowich1,3, Stephen M. Gordon1,2,
Chou P. Hung1,4, and Brent J. Lance1
1Human Research and Engineering Directorate, U.S. Army Research Laboratory, Aberdeen
Proving Ground, MD
2DCS Corporation, Alexandria, VA
3Department of Biomedical Engineering, Columbia University, New York, NY
4Department of Neuroscience, Georgetown University, Washington, DC
*Corresponding Author
May 17, 2018
Objective: Brain computer interfaces (BCI) enable direct communication with a computer,
using neural activity as the control signal. This neural signal is generally chosen from a va-
riety of well-studied electroencephalogram (EEG) signals. For a given BCI paradigm, feature
extractors and classifiers are tailored to the distinct characteristics of its expected EEG control
signal, limiting its application to that specific signal. Convolutional Neural Networks (CNNs),
which have been used in computer vision and speech recognition to perform automatic feature
extraction and classification, have successfully been applied to EEG-based BCIs; however, they
have mainly been applied to single BCI paradigms and thus it remains unclear how these archi-
tectures generalize to other paradigms. Here, we ask if we can design a single CNN architecture
to accurately classify EEG signals from different BCI paradigms, while simultaneously being as
compact as possible (defined as the number of parameters in the model). Approach: In this work
we introduce EEGNet, a compact convolutional neural network for EEG-based BCIs. We intro-
duce the use of depthwise and separable convolutions to construct an EEG-specific model which
encapsulates well-known EEG feature extraction concepts for BCI. We compare EEGNet, both
for within-subject and cross-subject classification, to current state-of-the-art approaches across
four BCI paradigms: P300 visual-evoked potentials, error-related negativity responses (ERN),
movement-related cortical potentials (MRCP), and sensory motor rhythms (SMR). Results: We
show that EEGNet generalizes across paradigms better than, and achieves comparably high
performance to, the reference algorithms when only limited training data is available. We also
show that EEGNet effectively generalizes to both ERP and oscillatory-based BCIs. In addition,
we demonstrate three different approaches to visualize the contents of a trained EEGNet model
to enable interpretation of the learned features. Significance: Our results suggest that EEGNet
is robust enough to learn a wide variety of interpretable features over a range of BCI tasks,
suggesting that the observed performances were not due to artifact or noise sources in the data.
Our models can be found at:
arXiv:1611.08024v4 [cs.LG] 16 May 2018
Keywords: Brain-Computer Interface, EEG, Deep Learning, Convolutional Neural Network,
P300, Error-Related Negativity, Sensory Motor Rhythm
1 Introduction
A Brain-Computer Interface (BCI) enables direct communication with a machine via brain sig-
nals [1]. Traditionally, BCIs have been used for medical applications such as neural control of
prosthetic artificial limbs [2]. However, recent research has opened up the possibility for novel BCIs
focused on enhancing performance of healthy users, often with noninvasive approaches based on
electroencephalography (EEG) [3–5]. Generally speaking, a BCI consists of five main processing
stages [6]: a data collection stage, where neural data is recorded; a signal processing stage, where
the recorded data is preprocessed and cleaned; a feature extraction stage, where meaningful infor-
mation is extracted from the neural data; a classification stage, where a decision is interpreted from
the data; and a feedback stage, where the result of that decision is provided to the user. While these
stages are largely the same across BCI paradigms, each paradigm relies on manual specification
of signal processing [7], feature extraction [8] and classification methods [9], a process which often
requires significant subject-matter expertise and/or a priori knowledge about the expected EEG
signal. It is also possible that, because the EEG signal preprocessing steps are often very specific
to the EEG feature of interest (for example, band-pass filtering to a specific frequency range), that
other potentially relevant EEG features could be excluded from analysis (for example, features
outside of the band-pass frequency range). The need for robust feature extraction techniques will
only continue to increase as BCI technologies evolve into new application domains [3–5,10–12].
Deep Learning has largely alleviated the need for manual feature extraction, achieving state-of-
the-art performance in fields such as computer vision and speech recognition [13, 14]. Specifically,
the use of deep convolutional neural networks (CNNs) has grown due in part to their success in
many challenging image classification problems [15–19], surpassing methods relying on hand-crafted
features (see [14] and [20] for recent reviews). Although the majority of BCI systems still rely on
the use of handcrafted features, many recent works have explored the application of Deep Learning
to EEG signals. For example, CNNs have been used for epilepsy prediction and monitoring [21–25],
for auditory music retrieval [26, 27], for detection of visual-evoked responses [28–31] and for motor
imagery classification [32], while Deep Belief Networks (DBNs) have been used for sleep stage
detection [33], anomaly detection [34] and in motion-onset visual-evoked potential classification [35].
CNNs using time-frequency transforms of EEG data were used for mental workload classification
[36] and for motor imagery classification [37–39]). Restricted Boltzman Machines (RBMs) have been
used for motor imagery [40]. An adaptive method based on stacked denoising autoencoders has been
proposed for mental workload classification [41]). These studies focused primarily on classification
in a single BCI task, often times using task-specific knowledge in designing the network architecture.
In addition, the amount of data used to train these networks varied significantly across studies, in
part due to the difficulty in collecting data under different experimental designs. Thus, it remains
unclear how these previous deep learning approaches would generalize both to other BCI tasks as
well as to variable training data sizes.
In this work we introduce EEGNet, a compact CNN for classification and interpretation of
EEG-based BCIs. We introduce the use of Depthwise and Separable convolutions, previously used
in computer vision [42], to construct an EEG-specific network that encapsulates several well-known
EEG feature extraction concepts, such as optimal spatial filtering and filter-bank construction,
while simultaneously reducing the number of trainable parameters to fit when compared to exist-
ing approaches. We evaluate the generalizability of EEGNet on EEG datasets collected from four
different BCI paradigms: P300 visual-evoked potential (P300), error-related negativity (ERN),
movement-related cortical potential (MRCP) and the sensory motor rhythm (SMR), representing
a spectrum of paradigms based on classification of Event-Related Potentials (P300, ERN, MRCP)
as well as classification of oscillatory components (SMR). In addition, each of these data collec-
tions contained varying amounts of data, allowing us to explore the efficacy of EEGNet on various
training data sizes. Our results are as follows: We show that EEGNet achieves improved classifi-
cation performance over an existing paradigm-agnostic EEG CNN model across nearly all tested
paradigms when limited training data is available. In addition, we show that EEGNet effectively
generalizes across all tested paradigms. We also show that EEGNet performs just as well as a
more paradigm-specific EEG CNN model, but with two orders of magnitude fewer parameters to
fit, representing a more efficient use of model parameters (an aspect that has been explored in
previous deep learning literature, see [42,43]). Finally, through the use of feature visualization and
model ablation analysis, we show that neurophysiologically interpretable features can be extracted
from the EEGNet model. This is important as CNNs, despite their ability for robust and auto-
matic feature extraction, often produce hard to interpret features. For neuroscience practitioners,
the ability to derive insights into CNN-derived neurophysiological phenomena may be just as im-
portant as achieving good classification performance, depending on the intended application. We
validate our architecture’s ability to extract neurophysiologically interpretable signals on several
well-studied BCI paradigms to show that the network performance is not being driven by noise or
artifact signals in the data.
The remainder of this manuscript is structured as follows. Section 2.1 gives a brief description
of the four datasets used to validate our CNN model. Section 2.2 describes our EEGNet model as
well as other BCI models (both CNN and non-CNN based models) used in our model comparison.
Section 3 presents the results of both within-subject and cross-subject classification performance,
as well as results of our feature explainability analysis. We discuss our findings in more detail in
the Discussion.
2 Materials and Methods
2.1 Data Description
BCIs are generally categorized into two types, depending on the EEG feature of interest [44]:
event-related and oscillatory. Event-Related Potential (ERP) BCIs are designed to detect a high
amplitude and low frequency EEG response to a known, time-locked external stimulus. They are
generally robust across subjects and contain well-stereotyped waveforms, enabling the time course
of the ERP to be modeled through machine learning efficiently [45]. In contrast to ERP-based BCIs,
which rely mainly on the detection of the ERP waveform from some external event or stimulus,
Paradigm Feature Type Bandpass Filter # of Subjects Trials per Subject # of Classes Class Imbalance?
P300 ERP 1-40Hz 15 2000 2 Yes, 5.6:1
ERN ERP 1-40Hz 26 340 2 Yes, 3.4:1
MRCP ERP/Oscillatory 0.1-40Hz 13 1100 2 No
SMR Oscillatory 4-40Hz 9 288 4 No
Table 1: Summary of the data collections used in this study. Class imbalance, if present, is given as
odds; i.e.: an odds of 2:1 means the class imbalance is 2/3 of the data for class 1 to 1/3 of the data
for class 2. For the P300 and ERN datasets, the class imbalance is subject-dependent; therefore,
the odds is given as the average class imbalance over all subjects.
Oscillatory BCIs use the signal power of specific EEG frequency bands for external control and are
generally asynchronous [46]. When oscillatory signals are time-locked to an external stimulus, they
can be represented through event-related spectral perturbation (ERSP) analyses [47]. Oscillatory
BCIs are more difficult to train, generally due to the lower signal-to-noise ratio (SNR) as well as
greater variation across subjects [46]. A summary of the data used in this manuscript can be found
in Table 1
2.1.1 Dataset 1: P300 Event-Related Potential (P300)
The P300 event-related potential is a stereotyped neural response to novel visual stimuli [48]. It is
commonly elicited with the visual oddball paradigm, where participants are shown repetitive “non-
target” visual stimuli that are interspersed with infrequent “target” stimuli at a fixed presentation
rate (for example, 1 Hz). Observed over the parietal cortex, the P300 waveform is a large positive
deflection of electrical activity observed approximately 300 ms post stimulus onset, the strength
of the observed deflection being inversely proportional to the frequency of the target stimuli. The
P300 ERP is one of the strongest neural signatures observable by EEG, especially when targets
are presented infrequently [48]. When the image presentation rate increases to 2 Hz or more, it is
commonly referred to as rapid serial visual presentation (RSVP), which has been used to develop
BCIs for large image database triage [49–51].
The EEG data used here have been previously described in [50]; a brief description is given
below. 18 participants volunteered for an RSVP BCI study. Participants were shown images of
natural scenery at 2 Hz rate, with images either containing a vehicle or person (target), or with no
vehicle or person present (non-target). Participants were instructed to press a button with their
dominant hand when a target image was shown. The target/non-target ratio was 20%/80%. Data
from 3 participants were excluded from the analysis due to excessive artifacts and/or noise within
the EEG data. Data from the remaining 15 participants (9 male and 14 right-handed) who ranged in
age from 18 to 57 years (mean age 39.5 years) were further analyzed. EEG recordings were digitally
sampled at 512 Hz from 64 scalp electrodes arranged in a 10-10 montage using a BioSemi Active
Two system (Amsterdam, The Netherlands). Continuous EEG data were referenced offline to the
average of the left and right earlobes, digitally bandpass filtered, using an FIR filter implemented
in EEGLAB [52], to 1-40 Hz and downsampled to 128 Hz. EEG trials of target and non-target
conditions were extracted at [0,1]spost stimulus onset, and used for a two-class classification.
2.1.2 Dataset 2: Feedback Error-Related Negativity (ERN)
Error-Related Negativity potentials are perturbations of the EEG following an erroneous or unusual
event in the subject’s environment or task. They can be observed in a variety of tasks, including time
interval production paradigms [53] and in forced-choice paradigms [54, 55]. Here we focus on the
feedback error-related negativity (ERN), which is an amplitude perturbation of the EEG following
the perception of an erroneous feedback produced by a BCI. The feedback ERN is characterized as
a negative error component approximately 350ms, followed by a positive component approximately
500ms, after visual feedback (see Figure 7 of [56] for an illustration). The detection of the feedback
ERN provides a mechanism to infer, and to possibly correct in real-time, the incorrect output of a
BCI. This two-stage system has been proposed as a hybrid BCI in [57,58] and has been shown to
improve the performance of a P300 speller in online applications [59].
The EEG data used here comes from [56] and was used in the “BCI Challenge” hosted by Kaggle
(; a brief description is given below. 26 healthy
participants (16 for training, 10 for testing) participated in a P300 speller task, a system which uses a
random sequence of flashing letters, arranged in a 6×6 grid, to elicit the P300 response [60]. The goal
of the challenge was to determine whether the feedback of the P300 speller was correct or incorrect.
The EEG data were originally recorded at 600Hz using 56 passive Ag/AgCl EEG sensors (VSM-
CTF compatible system) following the extended 10-20 system for electrode placement. Prior to our
analysis, the EEG data were band-pass filtered, using an FIR filter implemented in EEGLAB [52],
to 1-40 Hz and down-sampled to 128Hz. EEG trials of correct and incorrect feedback were extracted
at [0,1.25]spost feedback presentation and used as features for a two-class classification.
2.1.3 Dataset 3: Movement-Related Cortical Potential (MRCP)
Some neural activities contain both ERP as well as an oscillatory components. One particular
example of this is the movement-related cortical potential (MRCP), which can be elicited by vol-
untary movements of the hands and feet and is observable through EEG along the central and
midline electrodes, contralateral to the hand or foot movement [61–64]. The MRCP components
can be seen before movement onset (a slow 0-5Hz readiness potential [65,66] and an early desyn-
chronization in the 10-12Hz frequency band), at movement onset (a slow motor potential [66,67]),
and after movement onset (a late synchronization of 20-30Hz activity approximately 1 second after
movement execution). The MRCP has been used previously to develop motor control BCIs for
both healthy and physically disabled patients [68–70]
The EEG data used here have been previously described in [71]; a brief description is given
below. In this study, 13 subjects performed self-paced finger movements using the left index, left
middle, right index, or right middle fingers. The data was recorded using a 256 channel BioSemi
Active II system at 1024 Hz. Due to extensive signal noise present in the data, the EEG data
were first processed with the PREP pipeline [72]. The data were referenced to linked mastoids,
bandpass filtered, using an FIR filter implemented in EEGLAB [52], between 0.1 Hz and 40 Hz,
and then downsampled to 128 Hz. We further downsampled the channel space to the standard 64
channel BioSemi montage. The index and middle finger blocks for each hand were combined for
binary classification of movements originating from the left or right hand. EEG trials of left and
right hand finger movements were extracted at [.5,1]saround finger movement onset and used
for a two-class classification.
2.1.4 Dataset 4: Sensory Motor Rhythm (SMR)
A common control signal for oscillatory-based BCI is the sensorimotor rhythm (SMR), wherein
mu (8-12Hz) and beta (18-26Hz) bands desynchronize over the sensorimotor cortex contralateral
to an actual or imagined movement. The SMR is very similar to the oscillatory component of the
MRCP. Although SMR-based BCIs can facilitate nuanced, endogenous BCI control, they tend to
be weak and highly variable across and within subjects, conventionally demanding user-training
(neurofeedback) and long calibration times (20 minutes) in order to achieve reasonable performance
The EEG data used here comes from BCI Competition IV Dataset 2A [73] (called the SMR
dataset for the remainder of the manuscript). The data consists of four classes of imagined move-
ments of left and right hands, feet and tongue recorded from 9 subjects. The EEG data were
originally recorded using 22 Ag/AgCl electrodes, sampled at 250 Hz and bandpass filtered between
0.5 and 100Hz. We resampled the timeseries to 128 Hz, and follow the same EEG pre-processing
procedure as described in [32], using software that was provided by the authors. For both the
training and test sets we epoched the data at [0.5, 2.5] seconds post cue onset (the same window
which was used in [39,44]). Note that we make predictions for only this time range on the test set.
We perform a four-class classification using accuracy as the summary measure.
2.2 Classification Methods
2.2.1 EEGNet: Compact CNN Architecture
Here we introduce EEGNet, a compact CNN architecture for EEG-based BCIs that (1) can be
applied across several different BCI paradigms, (2) can be trained with very limited data and (3)
can produce neurophysiologically interpretable features. A visualization and full description of the
EEGNet model can be found in Figure 1 and Table 2, respectively, for EEG trials, collected at
128Hz sampling rate, having Cchannels and Ttime samples. We fit the model using the Adam
optimizer, using default parameters as described in [74], minimizing the categorical cross-entropy
loss function. We run 500 training iterations (epochs) and perform validation stopping, saving the
model weights which produced the lowest validation set loss. All models were trained on an NVIDIA
Quadro M6000 GPU, with CUDA 9 and cuDNN v7, in Tensorflow [75], using the Keras API [76].
We omit the use of bias units in all convolutional layers. Note that, while all convolutions are one-
dimensional, we use two-dimensional convolution functions for ease of software implementation.
Our software implementation can be found at
In Block 1, we perform two convolutional steps in sequence. First, we fit F12D convolutional
filters of size (1,64), with the filter length chosen to be half the sampling rate of the data
Figure 1: Overall visualization of the EEGNet architecture. Lines denote the convolutional kernel
connectivity between inputs and outputs (called feature maps) . The network starts with a temporal
convolution (second column) to learn frequency filters, then uses a depthwise convolution (middle
column), connected to each feature map individually, to learn frequency-specific spatial filters. The
separable convolution (fourth column) is a combination of a depthwise convolution, which learns a
temporal summary for each feature map individually, followed by a pointwise convolution, which
learns how to optimally mix the feature maps together. Full details about the network architecture
can be found in Table 2.
(here, 128Hz), outputting F1feature maps containing the EEG signal at different band-pass
frequencies. Setting the length of the temporal kernel at half the sampling rate allows for
capturing frequency information at 2Hz and above. We then use a Depthwise Convolution [42]
of size (C, 1) to learn a spatial filter. In CNN applications for computer vision the main
benefit of a depthwise convolution is reducing the number of trainable parameters to fit, as
these convolutions are not fully-connected to all previous feature maps (see Figure 1 for an
illustration). Importantly, when used in EEG-specific applications, this operation provides a
direct way to learn spatial filters for each temporal filter, thus enabling the efficient extraction
of frequency-specific spatial filters (see the middle column of Figure 1). A depth parameter
Dcontrols the number of spatial filters to learn for each feature map (D= 1 is shown in
Figure 1 for illustration purposes). This two-step convolutional sequence is inspired in part
by the Filter-Bank Common Spatial Pattern (FBCSP) algorithm [77] and is similar in nature
to another decomposition technique, Bilinear Discriminant Component Analysis [78]. We
keep both convolutions linear as we found no significant gains in performance when using
nonlinear activations. We apply Batch Normalization [79] along the feature map dimension
before applying the exponential linear unit (ELU) nonlinearity [80]. To help regularize or
model, we use the Dropout technique [81]. We set the dropout probability to 0.5 for within-
subject classification to help prevent over-fitting when training on small sample sizes, whereas
we set the dropout probability to 0.25 in cross-subject classification, as the training set sizes
Block Layer # filters size # params Output Activation Options
1 Input (C, T)
Reshape (1, C, T)
Conv2D F1(1, 64) 64 F1(F1, C, T) Linear mode = same
BatchNorm 2 F1(F1, C, T)
DepthwiseConv2D D * F1(C, 1) CDF1(D * F1, 1, T) Linear mode = valid, depth = D, max norm = 1
BatchNorm 2 DF1(D * F1, 1, T)
Activation (D * F1, 1, T) ELU
AveragePool2D (1, 4) (D * F1, 1, T // 4)
Dropout* (D * F1, 1, T // 4) p= 0.25 or p= 0.5
2 SeparableConv2D F2(1, 16) 16 DF1+F2(DF1) (F2, 1, T // 4) Linear mode = same
BatchNorm 2 F2(F2, 1, T // 4)
Activation (F2, 1, T // 4) ELU
AveragePool2D (1, 8) (F2, 1, T // 32)
Dropout* (F2, 1, T // 32) p= 0.25 or p= 0.5
Flatten (F2* (T // 32))
Classifier Dense N * (F2* T // 32) N Softmax max norm = 0.25
Table 2: EEGNet architecture, where C= number of channels, T= number of time points, F1=
number of temporal filters, D= depth multiplier (number of spatial filters), F2= number of
pointwise filters, and N= number of classes, respectively. For the Dropout layer, we use p= 0.5
for within-subject classification and p= 0.25 for cross-subject classification (see Section 2.1.1 for
more details)
are much larger (see Section 2.3 for more details on our within- and cross-subject analyses).
We apply an average pooling layer of size (1, 4) to reduce the sampling rate of the signal to
32Hz. We also regularize each spatial filter by using a maximum norm constraint of 1 on its
weights; kwk2<1.
In Block 2, we use a Separable Convolution, which is a Depthwise Convolution (here, of
size (1,16), representing 500ms of EEG activity at 32Hz) followed by F2(1,1) Pointwise
Convolutions [42]. The main benefits of separable convolutions are (1) reducing the number of
parameters to fit and (2) explicitly decoupling the relationship within and across feature maps
by first learning a kernel summarizing each feature map individually, then optimally merging
the outputs afterwards. When used for EEG-specific applications this operation separates
learning how to summarize individual feature maps in time (the depthwise convolution) with
how to optimally combine the feature maps (the pointwise convolution). This operation is also
particularly useful for EEG signals as different feature maps may represent data at different
time-scales of information. In our case we first learn a 500 ms “summary” of each feature
map, then combine the outputs afterwards. An Average Pooling layer of size (1,8) is used
for dimension reduction.
In the classification block, the features are passed directly to a softmax classification with N
units, Nbeing the number of classes in the data. We omit the use of a dense layer for feature
aggregation prior to the softmax classification layer to reduce the number of free parameters
in the model, inspired by the work in [82].
We investigate several different configurations of the EEGNet architecture by varying the num-
ber of filters, F1, and the number of spatial filters per temporal filter, Dto learn. We set F2=DF1
Trial Length (sec) DeepConvNet ShallowConvNet EEGNet-4,2 EEGNet-8,2
P300 1 174,127 104,002 1,066 2,258
ERN 1.25 169,927 91,602 1,082 2,290
MRCP 1.5 175,727 104,722 1,098 2,322
SMR* 2 152,219 40,644 796 1,716
Table 3: Number of trainable parameters per model and per dataset for all CNN-based models. We
see that the EEGNet models are up to two orders of magnitude smaller than both DeepConvNet
and ShallowConvNet across all datasets. Note that we use a temporal kernel length of 32 samples
for the SMR dataset as the data were high-passed at 4Hz.
(the number of temporal filters along with their associated spatial filters from Block 1) for the du-
ration of the manuscript, although in principle F2can take any value; F2< D F1denotes a
compressed representation, learning fewer feature maps than inputs, whereas F2> D F1denotes
an overcomplete representation, learning more feature maps than inputs. We use the notation
EEGNet-F1,D to denote the number of temporal and spatial filters to learn; i.e.: EEGNet-4,2
denotes learning 4 temporal filters and 2 spatial filters per temporal filter.
2.2.2 Comparison with existing CNN Approaches
We compare the performance of EEGNet against the DeepConvNet and ShallowConvNet models
proposed by [32]; full table descriptions of both models can be found in the Appendix. We imple-
mented these models in Tensorflow and Keras, following the descriptions found in the paper. As
their architectures were originally designed for 250Hz EEG signals (as opposed to 128Hz signals
used here) we divided the lengths of temporal kernels and pooling layers in their architectures by
2 to correspond approximately to the sampling rate used in our models. We train these models in
the same way we train the EEGNet model (see Section 2.2.1).
The DeepConvNet architecture consists of five convolutional layers with a softmax layer for
classification (see Figure 1 of [32]). The ShallowConvNet architecture consists of two convolutional
layers (temporal, then spatial), a squaring nonlinearity (f(x) = x2), an average pooling layer and
a log nonlinearity (f(x) = log(x)). We would like to emphasize that the ShallowConvNet archi-
tecture was designed specifically for oscillatory signal classification (by extracting features related
to log band-power); thus, it may not work well on ERP-based classification tasks. However, the
DeepConvNet architecture was designed to be a general-purpose architecture that is not restricted
to specific feature types [32], and thus it serves as a more valid comparison to EEGNet. Table 3
shows the number of trainable parameters per model across all CNN models.
2.2.3 Comparison with Traditional Approaches
We also compare the performance of EEGNet to that of the best performing traditional approach
for each individual paradigm. For all ERP-based data analyses (P300, ERN, MRCP) the tradi-
tional approach is the approach which won the Kaggle BCI Competition (code and documenta-
tion at, which uses a combination
of xDAWN Spatial Filtering [83], Riemannian Geometry [84,85], channel subset selection and L1
feature regularization (referred to as xDAWN + RG for the remainder of the manuscript). Here
we provide a summary of the approach, which is done in five steps:
1. Train two set of 5 xDAWN spatial filters, one set for each class of a binary classification task,
using the ERP template concatenation method as described in [85, 86].
2. Perform EEG electrode selection through backward elimination [87] to keep only the most
relevant 35 channels.
3. Project the covariance matrices onto the tangent space using the log-euclidean metric [84,88].
4. Perform feature normalization using an L1ratio of 0.5, signifying an equal weight for L1and
L2penalties. An L1penalty encourages the sum of the absolute values of the parameters to
be small, whereas an L2penalty encourages the sum of the squares of the parameters to be
small (a theoretical overview of these penalties can be found in [89]).
5. Perform classification using an Elastic Net regression.
We use the same xDAWN+RG model parameters across all comparisons (P300, ERN, MRCP)
with the exception of the initial number of EEG channels to use, which was set to 56 for ERN
and 64 for P300 and MRCP. While the original solution used an ensemble of bagged classifiers,
for this analysis we only compared a single model with this approach to a single EEGNet model
on identical training and test sets, as we expect any gains from ensemble learning to benefit both
approaches equally. The original solution also used a set of “meta features” that were specific to
that data collection. As the goal of this work is to investigate a general-purpose CNN model for
EEG-based BCIs, we omitted the use of these features as they are specific to that particular data
For oscillatory-based classification of SMR, the traditional approach is our own implementation
of the One-Versus-Rest (OVR) filter-bank common spatial pattern (FBCSP) algorithm as described
in [77]. Here we provide a brief summary of our approach:
1. Bandpass filter the EEG signal into 9 non-overlapping filter banks in 4Hz steps, starting at
4Hz: 4-8Hz, 8-12Hz, ..., 36-40Hz.
2. As the classification problem is multi-class, we use OVR classification, which requires that
we train a classifier for all pairs of OVR combinations, which there are 4 here (class 1 vs all
others, class 2 vs all others, etc). We train 2 CSP filter pairs (4 filters total) for each filter
bank on the training data using the auto-covariance shrinkage method by [90]. This will give
a total of 36 features (9 filter banks ×4 CSP filters) for each trial and each OVR combination.
3. Train an elastic-net logistic regression classifier [91] for each OVR combination. We set the
elastic net penalty α= 0.95.
4. Find the optimal λvalue for the elastic-net logistic regression that maximizes the validation
set accuracy by evaluating the trained classifiers on a held-out validation set. The multi-class
label for each trial is the classifier that produces the highest probability among the 4 OVR
5. Apply the trained classifiers to the test set, using the λvalues obtained in Step 4.
Note that this approach differs slightly from the original technique as proposed in [77], where they
use a Naive Bayes Parzen Window classifier. We opted to use an elastic net logistic regression for
ease of implementation, and the fact that it has been used in existing software implementations of
FBCSP (for example, in BCILAB [92]).
2.3 Data Analysis
Classification results are reported for two sets of analyses: within-subject and cross-subject. Within-
subject classification uses a portion of the subjects data to train a model specifically for that subject,
although cross-subject classification uses the data from other subjects to train a subject-agnostic
model. While within-subject models tend to perform better than cross-subject models on a variety
of tasks, there is ongoing research investigating techniques to minimize (or possibly eliminate) the
need for subject-specific information to train robust systems [44, 51].
For within-subject, we use four-fold blockwise cross-validation, where two of the four blocks
are chosen to be the training set, one block as the validation set, and the final block as testing.
We perform statistical testing using a repeated-measures Analysis of Variance (ANOVA), modeling
classification results (AUC for P300/MRCP/ERN and Classification Accuracy for SMR) as the
response variable with subject number and classifier type as factors. For cross-subject analysis in
P300 and MRCP we choose, at random, 4 subjects for the validation set, one subject for the test
set, and all remaining subjects for the training set (see Table 1 for number of subjects per dataset).
This process was repeated 30 times, producing 30 different folds. We follow the same procedure
for the ERN dataset, except we use the 10 test subjects from the original Kaggle Competition as
the test set for each fold. We perform statistical testing using a one-way Analysis of Variance,
using classifier type as the factor. For the SMR dataset, we partitioned the data as follows: For
each subject, select the training data from 5 other subjects at random to be the training set and
the training data from the remaining 3 subjects to be the validation set. The test set remains
the same as the original test set for the competition. Note that this enforces a fully cross-subject
classification analysis as we never use the test subjects’ training data. This process is repeated 10
times for each subject, creating 90 different folds. The mean and standard error of classification
performance were calculated over the 90 folds. We perform statistical testing for this analysis using
the same testing procedure as the within-subject analysis.
When training both the within-subject and cross-subject models, we apply a class-weight to
the loss function whenever the data is imbalanced (unequal number of trials for each class). The
class-weight we apply is the inverse of the proportion in the training data, with the majority class
set to 1. For example, in the P300 dataset, there is a 5.6:1 odds between non-targets and targets
(Table 1) . In this case the class-weight for non-targets was set to 1, while the class-weight for
targets was set to 6 (when the odds are a fraction, we take the next highest integer). This procedure
was applied to the P300 and ERN datasets only, as these were the only datasets where significant
class imbalance was present.
Note that for the SMR analysis, we set the temporal kernel length to be 32 samples long (as
opposed to 64 samples long as given in Table 2) since the data were high-passed at 4Hz.
2.4 EEGNet Feature Explainability
The development of methods for enabling feature explainability from deep neural networks has
become an active research area over the past few years, and has been proposed as an essential
component of a robust model validation procedure, to ensure that the classification performance
is being driven by relevant features as opposed to noise or artifacts in the data [16, 93–99]. We
present three different approaches for understanding the features derived by EEGNet:
1. Summarizing averaged outputs of hidden unit activations: This approach focuses
on summarizing the activations of hidden units at layers specified by the user. In this work
we choose to summarize the hidden unit activations representing the data after the depth-
wise convolution (the spatial filter operation in EEGNet). Because the spatial filters are tied
directly to a particular temporal filter, they provide additional insights into the spatial local-
ization of narrow-band frequency activity. Here we summarize the spatially-filtered data by
calculating the difference in averaged time-frequency representations between classes, using
Morlet wavelets [100].
2. Visualizing the convolutional kernel weights: This approach focuses on directly visual-
izing and interpreting the convolutional kernel weights from the model. Generally speaking,
interpreting the convolutional kernel weights is very difficult due to the cross-filter-map con-
nectivity between any two layers. However, because EEGNet limits the connectivity of the
convolutional layers (using depthwise and separable convolutions), it is possible to interpret
the temporal convolution as narrow-band frequency filters and the depthwise convolution as
frequency-specific spatial filters.
3. Calculating single-trial feature relevance on the classification decision: This ap-
proach focuses on calculating, on a single-trial basis, the relevance of individual features on
the resulting classification decision. Positive values of relevance denote evidence supporting
the outcome, while negative values of relevance denote evidence against the outcome. In
our analysis we used DeepLIFT with the Rescale rule [97], as implemented in [98], to calcu-
late single-trial EEG feature relevance. DeepLIFT is a gradient-based relevance attribution
method that calculates relevance values per feature relative to a “reference” input (here, an
input of zeros, as is suggested in [97]), and is a technique similar to Layerwise Relevance
Propagation (LRP) which has been used previously for EEG analysis [101] (a summary of
gradient-based relevance attribution methods can be found in [98]). This analysis can be used
to elucidate feature relevance from high-confidence versus low-confidence predictions, and can
4-fold Within-Subject Classification Performance
Figure 2: 4-fold within-subject classification performance for the P300, ERN and MRCP datasets
for each model, averaged over all folds and all subjects. Error bars denote 2 standard errors of
the mean. We see that, while there is minimal difference between all the CNN models for the
P300 dataset, there are significant differences in the MRCP dataset, with both EEGNet models
outperforming all other models. For the ERN dataset we also see both EEGNet models performing
better than all others (p < 0.05).
be used to confirm that the relevant features learned are interpretable, as opposed to noise
or artifact features.
3 Results
3.1 Within-Subject Classification
We compare the performance of both the CNN-based reference algorithms (DeepConvNet and
ShallowConvNet) and the traditional approach (xDAWN+RG for P300/MRCP/ERN and FBCSP
for SMR) with EEGNet-4,2 and EEGNet-8,2. Within-subject four-fold cross-validation results
across all algorithms for P300, MRCP and ERN datasets are shown in Figure 2. We observed,
across all paradigms, that there was no statistically significant difference between EEGNet-4,2
and EEGNet-8,2 (p > 0.05), indicating that the increase in model complexity did not statistically
improve classification performance. For the P300 dataset, all CNN-based models significantly out-
perform xDAWN+RG (p < 0.05) while not performing significantly different amongst themselves.
For the ERN dataset, EEGNet-8,2 outperforms DeepConvNet, ShallowConvNet and xDAWN+RG
(p < 0.05), while EEGNet-4,2 outperforms DeepConvNet and ShallowConvNet (p < 0.05). The
biggest difference observed among all the approaches is in the MRCP dataset, where both EEGNet
models statistically outperform all others by a significant margin (DeepConvNet, ShallowConvNet
and xDAWN+RG, p < 0.05 for each comparison).
Four-fold cross-validation results for the SMR dataset are shown in Figure 3. Here we see
the performances of ShallowConvNet and FBCSP are very similar, replicating previous results as
reported in [32], while DeepConvNet performance is significantly lower. We also see that EEGNet-
4-fold Within-Subject Classification Performance: SMR
Figure 3: 4-fold within-subject classification performance for the SMR dataset for each model,
averaged over all folds and all subjects. Error bars denote 2 standard errors of the mean. Here we
see DeepConvNet statistically performed worse than all other models (p < 0.05). ShallowConvNet
and EEGNet-8,2 performed similarly to that of FBCSP.
8,2 performance is similar to FBCSP as well.
3.2 Cross-Subject Classification
Cross-subject classification results across all algorithms for P300, MRCP and ERN datasets are
shown in Figure 4. Similar to the within-subject analysis, we observed no statistical difference
between EEGNet-4,2 and EEGNet-8,2 across all datasets (p > 0.05). For the P300 dataset, all
CNN-based models significantly outperform xDAWN+RG (p < 0.05) while not performing sig-
Cross-Subject Classification Performance
Figure 4: Cross-Subject classification performance for the P300, ERN and MRCP datasets for each
model, averaged for 30 folds. Error bars denote 2 standard errors of the mean. For the P300 and
MRCP datasets there is minimal difference between the DeepConvNet and the EEGNet models,
with both models outperforming ShallowConvNet. For the ERN dataset the reference algorithm
(xDAWN + RG) significantly outperforms all other models.
Cross-Subject Classification Performance: SMR
Figure 5: Cross-Subject classification performance for the SMR for each model, averaged over all
folds and all subjects. Error bars denote 2 standard errors of the mean. We see that all CNN-based
models perform similarly, while slightly outperforming FBCSP.
nificantly different amongst themselves. For the MRCP dataset EEGNet-8,2 and DeepConvNet
significantly outperform ShallowConvNet (p < 0.05). We also see that both DeepConvNet and
ShallowConvNet performance is better when compared to its within-subject performance for the
MRCP dataset. For the ERN dataset, xDAWN + RG outperforms all CNN models (p < 0.05).
Cross-subject classification results for the SMR dataset are shown in Figure 5, where we found no
significant difference in performance across all CNN-based models (p > 0.05).
3.3 EEGNet Feature Characterization
We illustrate three different approaches to characterize the features learned by EEGNet: (1) Sum-
marizing averaged outputs of hidden unit activations, (2) visualizing convolutional kernel weights,
and (3) calculating single-trial feature relevances on classification decision. We illustrate Approach
1 on the P300 dataset for a cross-subject trained EEGNet-4,1 model. We chose to analyze the filters
from the P300 dataset due to the fact that multiple neurophysiological events occur simultaneously:
participants were told to press a button with their dominant hand whenever a target image ap-
peared on the screen. Because of this, target trials contain both the P300 event-related potential
as well as the alpha/beta desynchronizations in contralateral motor cortex due to button presses.
Here we were interested in whether or not the EEGNet architecture was capable of separating out
these confounding events. We were also interested in quantifying the classification performance of
the architecture whenever specific filters were removed from the model.
Figure 6 shows the spatial topographies of the four filters along with an average wavelet time-
frequency difference, calculated using Morlet wavelets [100], between all target trials and all non-
target trials. Here we see four distinct filters appear. The time-frequency analysis of Filter 1 shows
an increase in low-frequency power approximately 500ms after image presentation, followed by
desynchronizations in alpha frequency. As nearly all subjects in the P300 dataset are right-handed,
we also see significant activity along the left motor cortex. Time-frequency analysis of Filter 2
Figure 6: Visualization of the features derived from an EEGNet-4,1 model configuration for one
particular cross-subject fold in the P300 dataset. (A) Spatial topoplots for each spatial filter. (B)
The mean wavelet time-frequency difference between target and non-target trials for each individual
appears to show a significant theta-beta relationship; while increases in theta activity have been
previously noted in the P300 literature in response to targets [102], a relationship between theta
and beta has not previously been noted. The time-frequency difference for Filter 4 appears to
correspond with the P300, with an increase low-frequency power approximately 350ms after image
Filters Removed Test Set AUC
(1) 0.8866
(2) 0.9076
(3) 0.8910
(4) 0.8747
(1, 2) 0.8875
(1, 3) 0.8593
(1, 4) 0.8325
(2, 3) 0.8923
(2, 4) 0.8721
(3, 4) 0.8206
(1, 2, 3) 0.8637
(1, 2, 4) 0.8202
(1, 3, 4) 0.7108
(2, 3, 4) 0.7970
None 0.9054
Table 4: Performance of a cross-subject trained EEGNet-4,1 model when removing certain filters
from the model, then using the model to predict the test set for one randomly chosen fold of the
P300 dataset. AUC values in bold denote the best performing model when removing 1, 2 or 3 filters
at a time. As the number of filters removed increases, we see decreases in classification performance,
although the magnitude of the decrease depends on which filters are removed.
Spat. Filter 1
Spat. Filter 2
Figure 7: Visualization of the features derived from a within-subject trained EEGNet-8,2 model for
Subject 3 of the SMR dataset. Each of the 8 columns shows the learned temporal kernel for a 0.25
second window (top) with its two associated spatial filters (bottom two). We see that, while many
of the temporal filters are isolating slower-wave activity, the network identifies a higher-frequency
filter at approximately 32Hz (Temp. Filter 3, which shows 8 cycles in a 0.25 s window).
We also conducted a feature ablation study, where we iteratively removed a set of filters (by
replacing the filters with zeros) and re-applied the model to predict trials in the test set. We do this
for all combinations of the four filters. Classification results for this ablation study are shown in
Table 4. We see that test set performance is minimally impacted by the removal of any single filter,
with the largest decrease occurring when removing Filter 4. As expected, when removing pairs of
filters the decrease in performance is more pronounced, with the largest decrease observed when
removing Filters 3 and 4. Removing Filters 2 and 3 results in practically no change in classification
performance when compared to the full model, suggesting that the most important features in
this task are being captured by Filters 1 and 4. This finding is further reinforced when looking
at classification performance when three filters are removed; a model that contains only Filter 4
(0.8637 AUC) performs fairly well when compared to models that contain only Filter 2 (0.7108
AUC) or Filter 1 (0.7970 AUC).
Figure 7 shows the filters learned for the EEGNet-8,2 model for a within-subject classification
of Subject 3 for the SMR dataset. Each column of this figure denotes the learned temporal kernel
(top row) with its two associated spatial filters (bottom two rows). Note that we are learning
temporal filters of length 32 samples, which correspond to 0.25 seconds in time; hence, we estimate
the frequency for each temporal filter as four times the number of observed cycles. Here we see that
EEGNet-8,2 learns both slow-frequency activity at approximately 12Hz (Filters 1, 2, 6 and 8, which
show three cycles in a 0.25 s window) and high-frequency activity at approximately 32Hz (Filter 3,
which show 8 cycles). Figure 8 compares the spatial filters associated with 8-12Hz frequency band
learned by EEGNet-8,2 with the spatial filters learned by FBCSP in the 8-12Hz filter-bank for each
of the four OVR combinations. For ease of description we will use the notation X-Y to denote the
row-column filter. Here we see many of the filters are strongly positively correlated across models
(i.e.: the 1-1 filter of EEGNet-8,2 with the 3-1 filter for FBCSP (ρ= 0.93) and the 2-1 filter of
EEGNet-8,2 with the 3-4 filter of FBCSP (ρ= 0.83)), while some are strongly negatively correlated
(the 3-1 filter of EEGNet-8,2 with the 1-1 filter of FBCSP (ρ=0.93)), indicating a similar filter
up to a sign ambiguity. This suggests that EEGNet, through the use of depthwise convolutions, is
Spatial Filter 1 Spatial Filter 2 Spatial Filter 3 Spatial Filter 4 Spatial Filter 1 Spatial Filter 2
Left hand vs. all Temporal Filter 1
Right hand vs. all Temporal Filter 2
Both feet vs. all Temporal Filter 6
Tongue vs. all Temporal Filter 8
FBCSP 8-12Hz Spatial Filters EEGNet-8,2 12Hz Spatial Filters
Figure 8: Comparison of the 4 spatial filters learned by FBCSP in the 8-12Hz filter bank for each
OVR class combination (A) with the spatial filters learned by EEGNet-8,2 (B) for 4 temporal filters
that capture 12Hz frequency activity for Subject 3 of the SMR dataset (Temporal Filters 1, 2, 6
and 8, see Figure 7). We see that similar filters appear across both FBCSP and EEGNet-8,2.
capable of learning band-specific spatial filters in a similar manner as FBCSP.
Figure 9 shows the single-trial feature relevances for EEGNet-8,2, calculated using DeepLIFT,
for three three different test trials for one cross-subject fold of the MRCP dataset. Here we see
that the high-confidence predictions (Figure 9A and Figure 9B, for left and right finger movement,
respectively) both correctly show the contralateral motor cortex relevance as expected, whereas for
a low-confidence prediction (Figure 9C), the feature relevance is more broadly distributed, both in
time and in space on the scalp.
Figure 10 shows an additional example of using DeepLIFT to analyze feature relevance for
a cross-subject trained EEGNet-4,2 model for one test subject of the ERN dataset. Margaux
et. al. [56] previously noted that the average ERP for correct feedback trials has an earlier peak
positive potential, corresponding to approximately 325 ms, whereas the positive peak potential
for incorrect trials occurs slightly later, approximately 475 ms. Here we see the same temporal
difference in the timing of the peak positive potential for incorrect feedback trials (vertical line in
top row of Figure 10) and correct feedback trials (vertical line in bottom row of Figure 10). We also
see the DeepLIFT feature relevances align very closely to that of the peak positive potential for
both classes, suggesting that the network has focused on the peak positive potential as the relevant
feature for ERN classification. This finding supports results previously reported in [56], where they
showed a strong positive correlation between the amplitude of the peak positive potential and the
accuracy of error detection.
Figure 9: (Top row) Single-trial EEG feature relevance for a cross-subject trained EEGNet-8,2
model, using DeepLIFT, for three different test trials of the MRCP dataset: (A) a high-confidence,
correct prediction of left finger movement, (B) a high-confidence, correct prediction of right finger
movement and (C) a low-confidence, incorrect prediction of left finger movement. Titles include the
true class label and the predicted probability of that label. (Bottom row) Spatial topoplots of the
relevances at two time points: approximately 50 ms and 150 ms after button press. As expected,
the high-confidence trials show the correct relevances corresponding to contralateral motor cortex
for left (A) and right (B) button presses, respectively. For the low-confidence trial we see the
relevances are more mixed and broadly distributed, without a clear spatial localization to motor
4 Discussion
In this work we proposed EEGNet, a compact convolutional neural network for EEG-based BCIs
that can generalize across different BCI paradigms in the presence of limited data and can produce
interpretable features. We evaluated EEGNet against the state-of-the-art approach for both ERP
and Oscillatory-based BCIs across four EEG datasets: P300 visual-evoked potentials, Error-Related
Negativity (ERN), Movement-Related Cortical Potentials (MRCP) and Sensory Motor Rhythms
(SMR). To the best of our knowledge, this represents the first work that has validated the use of a
single network architecture across multiple BCI datasets, each with their own feature characteristics
and data set sizes. Our work introduced the use of Depthwise and Separable Convolutions [42]
for EEG signal classification, and showed that they can be used to construct an EEG-specific
model which encapsulates well-known EEG feature extraction concepts. Finally, through the use of
feature visualization and ablation analysis, we show that neurophysiologically interpretable features
can be extracted from the EEGNet model, providing further validation and evidence that the
network performance is not being driven by noise or artifact signals in the data. This last finding is
particularly important, as it is a critical component to understanding the validity and robustness of
CNN model architectures not just for EEG [32,101], but for CNN architectures in general [16,94,99].
The learning capacity of CNNs comes in part from their ability to automatically extract intricate
feature representations from raw data. However, since the features are not hand-designed by
Figure 10: Single-trial EEG feature relevance for a cross-subject trained EEGNet-4,2 model, using
DeepLIFT, for the one test subject of the ERN dataset. (Top Row) Feature relevances for three
correctly predicted trials of incorrect feedback, along with its predicted probability P. (Bottom
Row) Same as the top row but for three correctly predicted trials of correct feedback. The black
line denotes the average ERP, calculated at channel Cz, for incorrect feedback trials (top row) and
for correct feedback trials (bottom row). The thin vertical line denotes the positive peak of the
average ERP waveform. Here we see feature relevances coincide strongly with the positive peak of
the average ERP waveform for each trial. We also see the positive peak occurring slightly earlier
for correct feedback trials versus incorrect feedback trials, consistent with the results in [56].
human engineers, understanding the meaning of those features poses a significant challenge in
producing interpretable models [95]. This is especially true when CNNs are used for the analysis
of EEG data where features from neural signals are often non-stationary and corrupted by noise
artifacts [103, 104]. In this study, we illustrated three different approaches for visualizing the
features learned by EEGNet: (1) analyzing spatial filter outputs, averaged over trials, on the P300
dataset, (2) visualizing the convolutional kernel weights on the SMR dataset and comparing them
to the weights learned by FBCSP, and (3) performing single-trial relevance analysis on the MRCP
and SMR datasets. For the ERN dataset we compared single-trial feature relevances to averaged
ERPs and showed that relevant features coincided with the peak of the positive potential for
correct and incorrect feedback trials, which has been shown in previous literature to be positively
correlated to classifier performance [56]. In addition, we conducted a feature ablation study to
understand the impact of a classification decision on the presence or absence of a particular feature
on the P300 dataset. In each of these analyses, we showed that EEGNet was capable of extracting
interpretable features that generally corresponded to known neurophysiological phenomena. These
results suggest that the classification performances we observed were not due to artifact or noise
sources in the data.
Our results showed that the spatial filters learned by EEGNet for temporal kernels around
12Hz were significantly correlated to the spatial filters learned by FBCSP in the 8-12Hz filter
bank for the SMR dataset. This is interesting to note, as the optimization criterion for CSP
(optimal variance separation) is different than the optimization criterion for EEGNet (minimum
cross-entropy loss). Because of this, it is not guaranteed that the learned filters from these methods
would be comparable. It was encouraging to see that many of the filters did in fact overlap (up to
a sign ambiguity), suggesting that EEGNet is learning a similar feature representation to that of
FBCSP. This analysis is directly enabled by EEGNet’s use of depthwise convolutions to tie spatial
filters directly to a temporal filter, an aspect that is unique to this model.
Generally speaking, the classification performance of DeepConvNet and EEGNet were similar
across all cross-subject analyses, whereas DeepConvNet performance was lower across nearly all
within-subject analyses (with the exception of P300). One possible explanation for this discrepancy
is the amount of training data used to train the model; in cross-subject analyses the training set
sizes were about 10-15 times larger than that of within-subject analyses. This suggests that Deep-
ConvNet is more data-intensive compared to EEGNet, an unsurprising result given that the model
size of DeepConvNet is two orders of magnitude larger than EEGNet (see Table 3). We believe this
intuition is consistent with the findings originally reported by the developers of DeepConvNet [32],
where they state that a training data augmentation strategy was needed to obtain good classifica-
tion performance on the SMR dataset. In contrast to their work, we show that EEGNet performed
well across all tested datasets without the need for data augmentation, making the model simpler
to use in practice.
In general we found that, both in within- and cross-subject analyses, that ShallowConvNet
tended to perform worse on the ERP BCI datasets than on the oscillatory BCI dataset (SMR),
while the opposite behavior was observed with DeepConvNet. We believe this is due to the fact
that the ShallowConvNet architecture was designed specifically to extract log bandpower features;
in situations where the dominant feature is signal amplitude (as is the case in many ERP BCIs),
ShallowConvNet performance tended to suffer. The opposite situation occurred with DeepConvNet;
as its architecture was not designed to extract frequency features, its performance was lower in
situations where frequency power is the dominant feature. In contrast, we found that EEGNet
performed just as well as ShallowConvNet in SMR classification and just as well as DeepConvNet
in ERP classification (and outperforming in the case of within-subject MRCP, ERN and SMR
classifications), suggesting that EEGNet is robust enough to learn a wide variety of features over a
range of BCI tasks.
The severe underperformance of ShallowConvNet on within-subject MRCP classification was
unexpected, given the similarity in neural responses between the MRCP and SMR, and the fact
that ShallowConvNet performed well on SMR. This discrepancy in performance is not due to the
amount of training data used, as within-subject MRCP classification has approximately 700 training
trials, evenly split among left and right finger movements, whereas the SMR dataset has only 192
training trials, evenly split among four classes. In addition, we did not observe large deviations
in ShallowConvNet performance on the other datasets (P300 and ERN). In fact, ShallowConvNet
performed fairly well on within-subject ERN classification, even though this dataset is the smallest
among all datasets used in this study (only having 170 training trials total). Determining the
underlying source of this phenomena will be explored in future research.
Deep Learning models for EEG generally employ one of three input styles, depending on their
targeted application: (1) the EEG signal of all available channels, (2) a transformed EEG signal
(generally a time-frequency decomposition) of all available channels [36] or (3) a transformed EEG
signal of a subset of channels [37]. Models that fall in (2) generally see a significant increase in data
dimensionality, thus requiring either more data or more model regularization (or both) to learn
an effective feature representation. This introduces more hyperparameters that must be learned,
increasing the potential variability in model performance due to hyperparameter misspecification.
Models that fall in (3) generally require a priori knowledge about the channels to select. For
example, the model proposed in [37] uses the time-frequency decomposition of channels Cz, C3
and C4 as the inputs for a motor imagery classification task. This channel selection is intentional,
given the fact that neural responses to motor actions (the sensory motor rhythm) are observed
strongest at those channels and are easily observed through a time-frequency analysis. Also, by
only working with three channels, the authors reduce the significant increase in dimensionality
of the data. While this approach works well if the feature of interest is known beforehand, this
approach is not guaranteed to work well in other applications where the features are not observed
at those channels, limiting the overall utility of this approach. We believe models that fall in (1),
such as EEGNet and others [28, 30, 31], offer the best tradeoff between input dimensionality and
the flexibility to discover relevant features by providing all available channels. This is especially
important as BCI technologies evolve into novel application spaces, as the features needed for these
future BCIs may not be known beforehand [3–5, 10–12].
This project was sponsored by the U.S. Army Research Laboratory under ARL-H70-HR52, ARL-
74A-HRCYB and through Cooperative Agreement Number W911NF-10-2-0022. The views and
conclusions contained in this document are those of the authors and should not be interpreted as
representing the official policies, either expressed or implied, of the U.S. Government. The U.S.
Government is authorized to reproduce and distribute reprints for Government purposes notwith-
standing any copyright notation herein.
Conflict of Interest Statement
The authors declare that the research was conducted in the absence of any commercial or financial
relationships that could be construed as a potential conflict of interest.
5 Appendix
5.1 DeepConvNet and ShallowConvNet architectures
The DeepConvNet and ShallowConvNet architectures are given in Tables 5 and 6, respectively. The
DeepConvNet was designed to be a general-purpose architecture that is not restricted to specific
feature types, whereas ShallowConvNet is designed specifically for oscillatory signal classification.
Layer # filters size # params Activation Options
Input (C, T)
Reshape (1, C, T)
Conv2D 25 (1, 5) 150 Linear mode = valid, max norm = 2
Conv2D 25 (C, 1) 25 * 25 * C + 25 Linear mode = valid, max norm = 2
BatchNorm 2 * 25 epsilon = 1e-05, momentum = 0.1
Activation ELU
MaxPool2D (1, 2)
Dropout p = 0.5
Conv2D 50 (1, 5) 25 * 50 * C + 50 Linear mode = valid, max norm = 2
BatchNorm 2 * 50 epsilon = 1e-05, momentum = 0.1
Activation ELU
MaxPool2D (1, 2)
Dropout p = 0.5
Conv2D 100 (1, 5) 50 * 100 * C + 100 Linear mode = valid, max norm = 2
BatchNorm 2 * 100 epsilon = 1e-05, momentum = 0.1
Activation ELU
MaxPool2D (1, 2)
Dropout p = 0.5
Conv2D 200 (1, 5) 100 * 200 * C + 200 Linear mode = valid, max norm = 2
BatchNorm 2 * 200 epsilon = 1e-05, momentum = 0.1
Activation ELU
MaxPool2D (1, 2)
Dropout p = 0.5
Dense N softmax max norm = 0.5
Table 5: DeepConvNet architecture, where C= number of channels, T= number of time points
and N= number of classes, respectively.
Layer # filters size # params Activation Options
Input (C, T)
Reshape (1, C, T)
Conv2D 40 (1, 13) 560 Linear mode = same, max norm = 2
Conv2D 40 (C, 1) 40 * 40 * C Linear mode = valid, max norm = 2
BatchNorm 2 * 40 epsilon = 1e-05, momentum = 0.1
Activation square
AveragePool2D (1, 35), stride (1, 7)
Activation log
Dropout p = 0.5
Dense N softmax max norm = 0.5
Table 6: ShallowConvNet architecture, where C= number of channels, T= number of time points
and N= number of classes, respectively. Here, the ’square’ and ’log’ activation functions are given
as f(x) = x2and f(x) = log(x), respectively. Note that we clip the log function such that the
minimum input value is a very small number (= 10e7) for numerical stability.
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... As such, CNNs appear suitable for analyzing long-term, multi-dimensional, and highly non-stationary signals, such as EEG signals. Several studies have explored the application of CNNs to EEG analysis in healthy populations in areas such as sensory processing, cognitive-emotional processing, speech, and motor planning/execution and have achieved excellent performance (Sussillo et al., 2016;Schirrmeister et al., 2017;Lawhern et al., 2018). ...
... A dense layer with softmax activation was used as a classification output layer, which divided the EEG segments into four levels. The network referred to previous research works to use CNN in EEG decoding, including, DeepConvNet, ShallowConvNet (Schirrmeister et al., 2017), and EEGNet (Lawhern et al., 2018). ...
... The network structure used here was in reference to previous research works to use CNN in EEG decoding, including, DeepConvNet, ShallowConvNet (Schirrmeister et al., 2017), and EEGNet (Lawhern et al., 2018). The results proved that the CNN-based model can achieve very high accuracy for BASED scoring. ...
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In recent years, the Burden of Amplitudes and Epileptiform Discharges (BASED) score has been used as a reliable, accurate, and feasible electroencephalogram (EEG) grading scale for infantile spasms. However, manual EEG annotation is, in general, very time-consuming, and BASED scoring is no exception. Convolutional neural networks (CNNs) have proven their great potential in many EEG classification problems. However, very few research studies have focused on the use of CNNs for BASED scoring, a challenging but vital task in the diagnosis and treatment of infantile spasms. This study proposes an automatic BASED scoring framework using EEG and a deep CNN. The feasibility of using CNN for automatic BASED scoring was investigated in 36 patients with infantile spasms by annotating their long-term EEG data with four levels of the BASED score (scores 5, 4, 3, and ≤2). In the validation set, the accuracy was 96.9% by applying a multi-layer CNN to classify the EEG data as a 4-label problem. The extensive experiments have demonstrated that our proposed approach offers high accuracy and, hence, is an important step toward an automatic BASED scoring algorithm. To the best of our knowledge, this is the first attempt to use a CNN to construct a BASED-based scoring model.
... CNN was the most widely used architecture for MI classification [2]. Standard CNN models with light [12] and deep architectures [19] have been proposed, as well as many other CNN varieties, including inception-based CNN [10], [11], residual-based CNN [20], 3D-CNN [20], multi-scale CNN [13], multi-layer CNN [18], multi-branch CNN [9], [20], and attention-based CNN [8]- [11], [13]. Several other DL models have also been suggested by some studies for classifying MI tasks. ...
... Some recent studies have used TCN architectures to classify MI tasks [22], [23]. Ingolfsson et al. [22] proposed a TCN model named EEG-TCN that combines TCN with the well-known EEGNet architecture [12]. A recent study in [23] attempted to improve the EEG-TCN model using the feature fusion technique. ...
... The CV block is similar to the EEGNet architecture proposed in [12]. CV block differs from EEGNet by using 2D convolution instead of separable convolution, which showed better performance. ...
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The brain-computer interface (BCI) is a cutting-edge technology that has the potential to change the world. Electroencephalogram (EEG) motor imagery (MI) signal has been used extensively in many BCI applications to assist disabled people, control devices or environments, and even augment human capabilities. However, the limited performance of brain signal decoding is restricting the broad growth of the BCI industry. In this paper, we propose an attention-based temporal convolutional network (ATCNet) for EEG-based motor imagery classification. The ATCNet model utilizes multiple techniques to boost the performance of MI classification with a relatively small number of parameters. ATCNet employs scientific machine learning to design a domain-specific DL model with interpretable and explainable features, multi-head self-attention to highlight the most valuable features in MI-EEG data, temporal convolutional network (TCN) to extract high-level temporal features, and convolutional-based sliding window to augment the MI-EEG data efficiently. The proposed model outperforms the current state-of-the-art techniques in the BCI Competition IV-2a dataset with an accuracy of 85.38% and 70.97% for the subject-dependent and subject-independent modes, respectively.
... It is still a worthy topic that how to design a practical deep learning framework to recognize and classify emotions from the original EEG signals directly. EEGNet (Lawhern et al., 2016) is a compact convolutional neural network suitable for EEG signals. Our study introduced extracting EEG features and classifying emotions by using depthwise separable convolution. ...
... Due to the solid internal relationship between different channels of EEG signal and the time correlations. Inspired by Lawhern et al. (2016), we proposed an end-to-end neural network (E2ENNet) for EEG emotion, which concatenates EEGNet and LSTM (Long-Short Term Memory). We use depthwise separable convolution to extract features from multi-channel original EEG signals, LSTM for searching the correlations between those features. ...
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Objectve Emotional brain-computer interface can recognize or regulate human emotions for workload detection and auxiliary diagnosis of mental illness. However, the existing EEG emotion recognition is carried out step by step in feature engineering and classification, resulting in high engineering complexity and limiting practical applications in traditional EEG emotion recognition tasks. We propose an end-to-end neural network, i.e., E2ENNet. Methods Baseline removal and sliding window slice used for preprocessing of the raw EEG signal, convolution blocks extracted features, LSTM network obtained the correlations of features, and the softmax function classified emotions. Results Extensive experiments in subject-dependent experimental protocol are conducted to evaluate the performance of the proposed E2ENNet, achieves state-of-the-art accuracy on three public datasets, i.e., 96.28% of 2-category experiment on DEAP dataset, 98.1% of 2-category experiment on DREAMER dataset, and 41.73% of 7-category experiment on MPED dataset. Conclusion Experimental results show that E2ENNet can directly extract more discriminative features from raw EEG signals. Significance This study provides a methodology for implementing a plug-and-play emotional brain-computer interface system.
... The Riemannian CSP was proposed to learn spatial features by replacing the Euclidean mean in the original CSP with the Riemannian mean (Barachant et al., 2010). Recently, deep learning methods with strong fitting ability for mass data have been developed for the rapid decoding of EEG signals using MI-BCI systems (Lawhern et al., 2018). Graph convolutional neural networks (GCNs-Net) was designed to filter EEG signals based on functional topological relationship to learn generalized features using graph convolutional layers . ...
... 4) GCNs-Net: A graph convolutional neural network for decoding motor imagery signals . 5) EEGNet: A compact convolutional neural network for EEG-based BCIs (Lawhern et al., 2018). ...
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Recently, motor imagery brain-computer interfaces (MI-BCIs) with stimulation systems have been developed in the field of motor function assistance and rehabilitation engineering. An efficient stimulation paradigm and Electroencephalogram (EEG) decoding method have been designed to enhance the performance of MI-BCI systems. Therefore, in this study, a multimodal dual-level stimulation paradigm is designed for lower-limb rehabilitation training, whereby visual and auditory stimulations act on the sensory organ while proprioceptive and functional electrical stimulations are provided to the lower limb. In addition, upper triangle filter bank sparse spatial pattern (UTFB-SSP) is proposed to automatically select the optimal frequency sub-bands related to desynchronization rhythm during enhanced imaginary movement to improve the decoding performance. The effectiveness of the proposed MI-BCI system is demonstrated on an the in-house experimental dataset and the BCI competition IV IIa dataset. The experimental results show that the proposed system can effectively enhance the MI performance by inducing the α, β and γ rhythms in lower-limb movement imagery tasks.
The motor imagery brain-computer interface (MI-BCI) based on electroencephalography (EEG) enables direct communication between the human brain and external devices. In this paper, the MTFB-CNN, a parallel multi-scale time-frequency block convolutional neural network based on the channel attention module, is proposed for EEG signals decoding, which can adaptively extract the time, frequency, and time-frequency domain features through parallel multi-scale time-frequency blocks, and then fuses and filters the features through attention mechanism and residual module. Experimental results based on the BCI Competition IV 2a and 2b datasets and the high gamma dataset show that the model achieves the highest average accuracy and kappa compared with existing baseline models. The MTFB-CNN is a novel and effective end-to-end model for decoding EEG signals without complex signals pre-processing operations, which has multi-scale feature extraction capability, making it successful in MI-BCI applications.
The early diagnosis of Alzheimer’s Disease (AD) plays a central role in the treatment of AD. Particularly, identifying the preclinical AD (pAD) stage could be crucial for timely treatment in the elderly. However, screening participants with pAD requires a series of psychological and neurological examinations. Thus, an efficient diagnostic tool is needed. Here, we recruited 91 elderly participants and collected 1 minute of resting-state electroencephalography data to classify participants as normal aging or diagnosed with pAD. We used deep neural networks (Deep ConvNet, EEGNet, EEG-TCNet, and cascade CRNN) in the within- and cross-subject paradigms for classification and found individual variations of classification accuracy in the cross-subject paradigm. Further, we proposed an individualized diagnostic strategy to identify neurophysiological similarities across participants and the proposed approach considering individual characteristics improved the diagnostic performance by approximately 20%. Our findings suggest that considering individual characteristics would be a breakthrough in diagnosing AD using deep neural networks.
The electroencephalogram (EEG) signal is commonly applied in the brain-computer interface (BCI) system of the motor imagery paradigm because it is noninvasive and has a high time resolution. This paper proposes a motor imagery classification method based on convolutional neural networks and Riemannian geometry to overcome the problem of noise and extreme values impacting motor imagery classification performance. The time-domain properties of EEG signals are extracted using multiscale temporal convolutions, whereas the spatial aspects of EEG signals are extracted using multiple convolutional kernels learned by spatial convolution. The extracted features are mapped to a Riemannian manifold space, and bilinear mapping and logarithmic operations are performed on the features to solve the problem of noise and extreme values. The effectiveness of the proposed method is validated using four types of motor imagery in the BCI competition IV dataset 2a to evaluate the classification ability. The experimental results show that the proposed approach has obvious advantages in the classification performance of motor imagery.
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This study proposed a novel attention-based temporal convolutional network (ATCNet) for EEG-based motor imagery classification that outperformed state-of-the-art techniques in MI-EEG classification using the BCI-2a dataset with an accuracy of 85.4% and 71% for the subject-dependent and subject-independent modes, respectively. These high results came with a relatively small number of parameters (115.2K), which makes ATCNet applicable to limited devices. The ablation analysis showed that each block in the ATCNet model made a significant contribution to the performance of the ATCNet model. The proposed model demonstrated a powerful ability to extract MI features from a raw EEG signal without pre-processing using a limited-size and challenging dataset.
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Recent work has shown that convolutional networks can be substantially deeper, more accurate and efficient to train if they contain shorter connections between layers close to the input and those close to the output. In this paper we embrace this observation and introduce the Dense Convolutional Network (DenseNet), where each layer is directly connected to every other layer in a feed-forward fashion. Whereas traditional convolutional networks with L layers have L connections, one between each layer and its subsequent layer (treating the input as layer 0), our network has L(L+1)/2 direct connections. For each layer, the feature maps of all preceding layers are treated as separate inputs whereas its own feature maps are passed on as inputs to all subsequent layers. Our proposed connectivity pattern has several compelling advantages: it alleviates the vanishing gradient problem and strengthens feature propagation; despite the increase in connections, it encourages feature reuse and leads to a substantial reduction of parameters; its models tend to generalize surprisingly well. We evaluate our proposed architecture on five highly competitive object recognition benchmark tasks. The DenseNet obtains significant improvements over the state-of-the-art on all five of them (e.g., yielding 3.74% test error on CIFAR-10, 19.25% on CIFAR-100 and 1.59% on SVHN).
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Deep learning with convolutional neural networks (deep ConvNets) has revolutionized computer vision through end-to-end learning, that is, learning from the raw data. There is increasing interest in using deep ConvNets for end-to-end EEG analysis, but a better understanding of how to design and train ConvNets for end-to-end EEG decoding and how to visualize the informative EEG features the ConvNets learn is still needed. Here, we studied deep ConvNets with a range of different architectures, designed for decoding imagined or executed tasks from raw EEG. Our results show that recent advances from the machine learning field, including batch normalization and exponential linear units, together with a cropped training strategy, boosted the deep ConvNets decoding performance, reaching at least as good performance as the widely used filter bank common spatial patterns (FBCSP) algorithm (mean decoding accuracies 82.1% FBCSP, 84.0% deep ConvNets). While FBCSP is designed to use spectral power modulations, the features used by ConvNets are not fixed a priori. Our novel methods for visualizing the learned features demonstrated that ConvNets indeed learned to use spectral power modulations in the alpha, beta, and high gamma frequencies, and proved useful for spatially mapping the learned features by revealing the topography of the causal contributions of features in different frequency bands to the decoding decision. Our study thus shows how to design and train ConvNets to decode task-related information from the raw EEG without handcrafted features and highlights the potential of deep ConvNets combined with advanced visualization techniques for EEG-based brain mapping. Hum Brain Mapp, 2017.
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The movement related cortical potential (MRCP), a slow cortical potential from the scalp electroencephalogram (EEG), has been used in real-time brain-computer-interface (BCI) systems designed for neurorehabilitation. Detecting MPCPs in real time with high accuracy and low latency is essential in these applications. In this study, we propose a new MRCP detection method based on constrained independent component analysis (cICA). The method was tested for MRCP detection during executed and imagined ankle dorsiflexion of 24 healthy participants, and compared with four commonly used spatial filters for MRCP detection in an offline experiment. The effect of cICA and the compared spatial filters on the morphology of the extracted MRCP was evaluated by two indices quantifying the signal-to-noise ratio and variability of the extracted MRCP. The performance of the filters for detection was then directly compared for accuracy and latency. The latency obtained with cICA (−34 ± 29 ms motor execution (ME) and 28 ± 16 ms for motor imagery (MI) dataset) was significantly smaller than with all other spatial filters. Moreover, cICA resulted in greater true positive rates (87.11 ± 11.73 for ME and 86.66 ± 6.96 for MI dataset) and lower false positive rates (20.69 ± 13.68 for ME and 19.31 ± 12.60 for MI dataset) compared to the other methods. These results confirm the superiority of cICA in MRCP detection with respect to previously proposed EEG filtering approaches.
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The purported "black box"' nature of neural networks is a barrier to adoption in applications where interpretability is essential. Here we present DeepLIFT (Deep Learning Important FeaTures), a method for decomposing the output prediction of a neural network on a specific input by backpropagating the contributions of all neurons in the network to every feature of the input. DeepLIFT compares the activation of each neuron to its 'reference activation' and assigns contribution scores according to the difference. By optionally giving separate consideration to positive and negative contributions, DeepLIFT can also reveal dependencies which are missed by other approaches. Scores can be computed efficiently in a single backward pass. We apply DeepLIFT to models trained on MNIST and simulated genomic data, and show significant advantages over gradient-based methods. A detailed video tutorial on the method is at and code is at
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