Wireless Brain Signal Recordings based on Capacitive Electrodes Mehrnaz Kh. Hazrati * Graduate School for Computing in Medicine and Life Sciences, Institute for Signal Processing University of Lübeck Lübeck, Germany [email protected]Haliza Mat Husin Institute for Signal Processing University of Lübeck Lübeck, Germany Ulrich G. Hofmann Neuroelectronic Systems, Department of Neurosurgery, University Medical Center Freiburg Freiburg, Germany [email protected]Abstract— In this paper the development of a wireless electroencephalogram (EEG) monitoring system is presented. The system is capable of processing brain signals on-board recorded from non-contact sensors. The non-contact sensor was designed utilizing capacitive coupling as recording interface. The on-board multi-channel signal processing is performed on a tiny computer module with low power consumption, high performance embedded computing platform that can communicate via WiFi or Bluetooth. It provides an excellent option for developing a compact Brain Computer Interface (BCI) with a direct connection to the external device e. g. robot or prosthesis without employing a personal computer (PC). Index Terms—Brain signal, electroencephalogram monitoring system, capacitive electrode, Overo Fire, Gumstix®. I. INTRODUCTION Electroencephalography (EEG) is one of the most widely used methods for evaluating the electrical activities of the brain. Due to the advantages of non-invasive measurement and the capability of long term monitoring of the EEG signal, an electroencephalograph plays an important role in brain examinations and studies. EEG is a standard procedure used in clinical and research applications, especially, in the diagnosis of brain diseases such as epilepsy, sleeping disorder and abnormal behavior [1]. EEG measurements can be divided into two types based on the manner in which electric current is collected from the tissue and potential differences are measured; (1) contact sensors (also called electrodes) are typically made of metal and need to come in direct electrical contact with the skin [10, 11]. This galvanic contact implies that there is a transfer of electric current between the sensors and the tissue. In that case electric currents are transferred through an electrically- conductive gel. Gelled elctrodes are often referred to as wet selectrodes, otherwise they are called dry electrodes [1, 11], (2) capacitive sensors are equipped with an insulated conducting plate placed in close contact with the body’s surface. This plate and the body’s surface form a capacitor and the electric signal can be derived over it. This is referred to as capacitive-coupling measurement [2, 5]. Two issues limit the use of commercially available EEG recording systems. The primary limiting issue is that they are hardly portable. Commercially available EEG recording systems are usually wired to a stationary computer, in order to save and to analyze the data. In addition recording electrodes are usually mounted on a tight cap that the subject wears on the head. The EEG cap is not portable due to the mass of wires connecting the cap to the data collecting system, which itself has a high power consumption. The second problem is that electrically-conductive gel is often required for a good connection between the mounted sensors and the scalp. The gel takes a lot of time to apply and it tends to dry out, which limits the recording time. Dry electrodes depend on ionic contact at the electrode-electrolyte interface as for the wet electrode type. The difference is that the skin supplies the electrolyte in the form of perspiration [2], rather than relying on an artificial electrolyte. The impedance of dry electrodes is much higher than that for wet electrodes, so a buffer must be used to convert high impedance to low impedance in order to minimize the effects of the various noise sources, which are the main issue in designing these electrodes. As an alternative, capacitively coupled electrodes have an advantage over galvanic contact electrodes: they can be used without electrolyte gel on the unprepared skin and are immune to voltage drifts that could appear because the electrode-skin resistance changes [2, 4, 8]. To tackle the mentioned issues, one approach is the use of capacitive electrodes to provide analog brain signal sensing [2] along with a small lightweight embedded computer for signal processing on the head. In this work, we introduce a wireless EEG monitoring system which is capable of processing brain signals on-board recorded from non-contact sensors. In the following the elements of the system are introduced and some recorded data are validated. II.DESIGN OF CAPACITIVE ELECTRODES Capacitive electrodes allow acquiring bio-potentials through displacement currents instead of real charge currents since the electrolyte-electrode-skin interface is replaced by a dielectric insulating film. Due to the absence of electrolyte, capacitive electrodes have a different behavior concerning the skin contact. A capacitive method of picking up biopotentials was proposed in
Electroencephalography (EEG) is one of the most widely
used methods for evaluating the electrical activities of the
brain. Due to the advantages of non-invasive measurement and
the capability of long term monitoring of the EEG signal, anelectroencephalograph plays an important role in brain
examinations and studies. EEG is a standard procedure used in
clinical and research applications, especially, in the diagnosis
of brain diseases such as epilepsy, sleeping disorder andabnormal behavior [1].
EEG measurements can be divided into two types based on
the manner in which electric current is collected from thetissue and potential differences are measured; (1) contact
sensors (also called electrodes) are typically made of metal
and need to come in direct electrical contact with the skin [10,
11]. This galvanic contact implies that there is a transfer of
electric current between the sensors and the tissue. In that case
electric currents are transferred through an electrically-
conductive gel. Gelled elctrodes are often referred to as wet
selectrodes, otherwise they are called dry electrodes [1, 11],
(2) capacitive sensors are equipped with an insulated
conducting plate placed in close contact with the body’ssurface. This plate and the body’s surface form a capacitor and
the electric signal can be derived over it. This is referred to ascapacitive-coupling measurement [2, 5].
Two issues limit the use of commercially available EEG
recording systems. The primary limiting issue is that they are
hardly portable. Commercially available EEG recording
systems are usually wired to a stationary computer, in order to
save and to analyze the data. In addition recording electrodes
are usually mounted on a tight cap that the subject wears onthe head. The EEG cap is not portable due to the mass of wires
connecting the cap to the data collecting system, which itself
has a high power consumption.
The second problem is that electrically-conductive gel is
often required for a good connection between the mounted
sensors and the scalp. The gel takes a lot of time to apply and it
tends to dry out, which limits the recording time. Dry
electrodes depend on ionic contact at the electrode-electrolyte
interface as for the wet electrode type. The difference is that
the skin supplies the electrolyte in the form of perspiration [2],
rather than relying on an artificial electrolyte. The impedance
of dry electrodes is much higher than that for wet electrodes, so
a buffer must be used to convert high impedance to low
impedance in order to minimize the effects of the various noise
sources, which are the main issue in designing these electrodes.
As an alternative, capacitively coupled electrodes have an advantage over galvanic contact electrodes: they can be used
without electrolyte gel on the unprepared skin and are immune
to voltage drifts that could appear because the electrode-skin
resistance changes [2, 4, 8].
To tackle the mentioned issues, one approach is the use of
capacitive electrodes to provide analog brain signal sensing [2]
along with a small lightweight embedded computer for signal processing on the head. In this work, we introduce a wirelessEEG monitoring system which is capable of processing brain
signals on-board recorded from non-contact sensors. In the
following the elements of the system are introduced and some
through displacement currents instead of real charge currents
since the electrolyte-electrode-skin interface is replaced by a dielectric insulating film. Due to the absence of electrolyte,
capacitive electrodes have a different behavior concerning theskin contact. A capacitive method of picking up biopotentials
was proposed in
1967 by Richardson [2] where the active electrode concept was
taken into account. The results demonstrated that capacitive
electrodes can be used to pick-up ECG signals with good
signal characteristics in comparison to wet electrodes. Up till
now, capacitive electrodes have mostly being studied for ECG
recordings [7, 13, 16, 18]. The new procedure based on capacitive measurements of skull potentials, while no direct
electrical contact with the scalp is made and therefore no gel is
needed, was first reported by Matsuo et al in 1973 using
barium titanate as the insulating material [3].
In 2004, recordings of ECG signals on a toilet seat have been performed by Kim et al. at the Advanced Biometric
Research Center, Seoul National University [19]. Their
capacitively-coupled electrode is composed of a Cu plate and a PTFE film. Kim et al. (2008) also made an important
contribution [12] by introducing a capacitively-coupled active
ground using and extending the driven-right-leg scheme described by Webster in 1983 [5]. In particular, they showed
that an active ground is highly effective at reducing line noise.
Researchers at Quantum Applied Science and Research
(QUASAR) developed a sensor that is able to measure the
ECG of a fully clothed person standing within a range of about
25 cm. In 2005, QUASAR developed a compact version of the
sensor and named it the capacitively coupled noncontact
electrode (CCNE), specifically to measure ECG through
clothing [16]. Sullivan et al. (2007) from Institute for Neural
Computation, University of California San Diego, designed an
integrated sensor which combines amplifications, band-pass
filtering, and analog-to-digital conversion within a 1 inch
diameter enclosure [9]. This non-contact bio-potential sensor
couples capacitively to the human scalp through hair for EEG
and to chest through clothing for ECG recordings [13, 18]. A solid copper fill forms a parallel plate capacitor with the body
and works as a capacitive electrode by sensing signals through
insulation such as fabric.
In 2008 at the “Institut für Elektrische Messtechnik und
Grundlagen der Elektrotechnik” of the TU Braunschweig,
Oehler et al. designed a capacitive electrode for EEG
measurements through hair [7]. 28 electrodes were integrated into an adjustable helmet to allow for different head shapes.
The state of the art of capacitive electrodes was reviewed in
the work of Spinelli and Haberman in 2010 [14]. They
designed capacitive electrodes that allow detection of bio-
potentials through thin clothes. On the other hand, they
reported that a sophisticated shielding and guarding at the
front-end stages is required to reduce high movement artifacts
and power line interferences. Recently Chi et al. designed an
innovative micro power non-contact EEG electrode with
active common-mode noise suppression and input capacitance
cancellation [17].
Our active electrode concept has adopted the capacitive
method by Richardson [2]. In contrast to the previous built
electrodes, we employed a combination of low noise, low
power electronic components such as an instrumentation amplifier LTC6079 and capacitance cancellation scheme.
III. SYSTEM DESIGN
The proposed EEG monitoring system is composed of two
main components: an analog unit running multi-channel
capacitive electrodes which include actively driven grounding,
and a digital unit based on a tiny computer module, Overo Fire
made by Gumstix® that has the capability to communicate via
WiFi and Bluetooth [22]. The system block diagram is shown
in Fig. 1 acquiring scalp potentials with the analog unit which in turn is connected to the digital unit responsible for analog to
digital conversion and signal processing.
Fig. 1. Block diagram of the wireless EEG recording system using Gumstix
commercial computer unit and Summit extension board
A. Analog Unit
The capacitive electrode was designed so that its input
impedance would be significantly larger than that of the skin-
electrode impedance to minimize interference caused by
motion artifact and unwanted common-mode voltages. The
signal on the skin capacitively couples to the sensing plate
coated with dielectric material to achieve the capacitive effect.
The coupling capacitance depends mainly on the thickness and
the dielectric constant of the material located between the
electrode and the subject’s skin [4]. The electrode surface
forms a coupling capacitance between the subject’s body and
the electrode as shown in Fig.2. The coupling capacitance depends mainly on the thickness and the dielectric constant, of
the dielectric material located between the electrode and the
subject’s skin. Using a capacitor model formula, the amount of
capacitive coupling,
present can be estimated, where is the relative static
permittivity of dielectric, is the permittivity of free space,
is the surface area of the plates and is the thickness of
dielectric.
Fig. 2. Capacitive sensing method
For our electrode, the signal on the skin capacitively
couples to the sensing plate at the bottom of the PC board,
which is covered with soldering mask for electrical isolation
of the sensor. The effective surface area of the 30 mm
diameter electrode board is A = π.r2 = 709 mm
2. To achieve
the capacitive effect, the sensing plate had to be coated with
dielectric. For simplicity, a cellophane tape based on cellulose
was chosen as dielectric with a dielectric constant of 3.9 [23] and the thickness of 0.058 mm. By using the capacitor model
equation, the approximate amount of capacitive coupling that
could be achieve by an electrode with diameter 30 mm is:
The general structure of a capacitive electrode is simplified
as in Fig. 3. The implementing circuit design behind a
capacitive electrode neccessarily needs to achieve low current
noise levels [4, 6]. Fig. 4 shows the top of our finished
electrode design and the connection between two units. The electronic circuit that is attached to the backside of the sensing
plate of the electrode incorporates a complex combination of
circuit techniques in order to achieve the required low current noise levels [5, 6]. It was built based on operational amplifiers
with very high input impedance in the order of TΩ to PΩ so the lower frequency limit of EEG signals is not compromised
[5]. The dominance of the resistive contribution to the total
input impedance requires an input biasing circuit to maintain a
stable dc operating point [5, 6]. Neutralization of input
capacitance is needed to ensure the electrode’s gain is constant
over a wide range of coupling distances thus making the
device applicable for diagnostic use [4]. Active shielding also
known as guarding, shields the electrode and electronics to
prevent the influence of disruptive external electric signals [7].
Fig. 3. General structure of capacitive electrode
The PCB design for the capacitive electrode consists of a 30 mm diameter four-layer PCB. The top layer houses the
component and traces, second layer is ground plate, third layer
is shielding plate and the bottom layer is the sensor plate. Such
PCB stacks are recommended for capacitive coupled sensing
applications [39, 40]. Fig. 4(a) presents each PCB layer of the
electrode, whereas Fig. 4 (b) illustrates the thickness of copper
and isolation for each of these layers.
(a)
(b)
Fig. 4. (a) Four-layer PCB of capacitive electrodes, layer 1 is for traces, layer
2 is the ground plane, layer 3 is the shielding plate and layer 4 is the
sensing plate, (b) Copper and Isolation thickness for PCB layers in
Eagle design software.
.
Number 1 and 16 represent the top and bottom layer
respectively and number 2 and 15 are the 2 layers in between.
Fig. 5 shows the top layer of the capacitive electrode board
from Eagle layout and the actual board. The reference DRL
electrode was also constructed of a four-layer PCB but with with 19 mm in diameter smaller in size compared to capacitive
electrodes. This is due to the less complex circuit required for
this electrode. To have a fully capacitive measurement system its bottom plate connects to the body capacitively like the
other electrodes.
B. Digital Unit
For the digital unit, we chose to develop on the Overo Fire,
a Linux-based computer-on-module (COM) produced by
Gumstix®. The Overo is a fully functional computer
motherboard that uses the Texas Instruments OMAP 3503
Application Processor. With roughly the size of a stick of gum,
this tiny board includes DSP (TMS320C64x) and GPU (ARM
Cortex-A8) as well as Bluetooth and the 802.11g wireless
networking protocols [21, 22]. The Overo Fire is an
appropriate embedded computing platform within our system‘s
requirements because of its minimal power requirements. As
consequence of its minute size and weight it can be placed on
patient’s head and still has an expansion port to add sensors.
The Summit is an expansion board that helps to support the
features of the Overo. It features a power supply for both itself
and the Overo [22]. The connection with the analog sensor
array is established via the ADC pin of the Summit board. The
10-bit ADC lines at the Overo board are controlled by the
TPS65950 Audio and Power Management module with a
maximum input voltage of 2.5V [21, 22].
(a)
(b)
Fig. 5. (a)Top view of the built capacitive electrode, (b) Analog and digital
units.
Since the Overo is a development platform, the appropriate
kernels to drive the analog sensors had to be configured, next
to programming the acquisition algorithms.
C. Gumstix ADC
Analog to Digital Converter (ADC) lines run direct from
the TPS65950 Power Management module in the Overo Fire
board with an input range of 0 to 2.5 V. ADC’s pin 2-7 can be accessed at the 40-pin header (SV1) at the Summit expansion
board [35,36]. The TPS65950 platform device supports
numerous functions which most of them are power related [9,
11]. It also supports the ADC module through twl4030-madc
driver and is connected to the OMAP chip on the I2C bus.
TWL3040 is a comparable older Texas Instruments chip which uses uses the same ADC module as the TPS65950 [35]. ADC driver can be accessed through a device input output control
(ioctl) interface, a single system call by which userspace can
communicate directly with the device driver [35, 37]. This
ioctl function was used to send control codes to the ADC.
There are sample programs provided by the developer to
access ADC driver depending on the version of Linux kernel
run on the Gumstix [22].
D. The Connector Board
The sensor connector board bridges the Summit and Overo
Fire board with sensors, and was also responsible for
supplying power to sensors. With the dimension of 80 mm x
39 mm, it is exactly the same size as the Summit board and allows the two boards to be be stacked together thus providing extra mechanical stability. EEG data was stored into an
external SD-memory card on the Gumstix. Finally the sensor
data was manually transferred to the host machine.
Connectivity between sensors and Gumstix was provided
by a flexible flat cable through a custom board connected with
a 40-pin header to the Summit board where the ADC pins are located. The ADC driver at the Overo kernel was interrupted
using a system call made at the Linux userspace. Fig.6 shows
the PCB of the connector board. The top layer of the board is
the signal routing of analog power supply, sensor input and
output, shielding and analog ground node for all electrodes.
The bottom layer is where all the low contact connectors
coming from electrodes are placed.
Fig. 6. The sensor connector board bridges the Summit and Overo Fire board
with sensors, and was also responsible for supplying power to sensors
II. RESULTS
Before designing the hardware, an AC sweep simulation
was applied in PSpice to evaluate the skin-electrode-analog
unit based on an existing skin-electrode model. The simulated
sensor circuit provides differential gain over a bandwidth of 1- 100Hz. In practice, any 50 Hz voltages from the power supply
line may cause a common mode voltage on the subject’s body.
Capacitive, actively driven ground connection using a well-
known technique by Webster [5] was applied for the purpose
of reducing the power line noise. We further applied a notch
filter at 50 Hz to minimize the effect of this noise in recorded
signals. Fig. 7 shows a time-domain signal of 8 seconds of
EEG data collected from Ag-AgCl and capacitive electrode at
Fp1 during an eyes closed, relax period. The red line
represents the EEG signal from capacitive electrodes, while the
blue line represents the EEG signal from the classical conductive electrode. The bottom window in each figure
represents a plot of one second signal extracted from the same
EEG data. Correlation coefficient between two signals during
eye closed state was calculated 0.7169 and this value was
0.6853 during eye open state.
Fig. 7. A comparison between designed capacitive electrode and a
commercial electrode: Recorded EEG of Fp2 while subject closed his
eyes, eight seconds samples (top) and one second samples (bottom).
Fig. 8 illustrates a 2min recorded signal from the occipital
position, when the subject closed his eyes for 60 sec in the middle of the recording. Note the increase in both amplitude
and α activity (around 8-10 Hz as seen in the zoom of the
spectrogram Fig. 6).
Fig. 8. Time series and spectrogram of 30-60-30 eye activities recorded with
the capacitive electrode. α activity can served be around 8-10 Hz.
III. CONCLUSION
An EEG monitoring system for non-contact sensing has
been presented in this paper. The use of capacitive coupling
technique in acquiring very low amplitude of brain signal
eliminates the need for electrolyte gel and skin abrasion.
Future work of the sensor will include the mechanical stability
of the electrode placement on the body to improve signal
quality as the next step towards a portable monitoring device.
Overall, we consider the combination of custom made analog
and off-the-shelve embedded system to provide the basis for
continuous brain monitoring for extended periods of time. The
final device will feature a compact design; low power
consumption as well as efficient data transmission and
processing due to its on-chip DSP and the option to directly
connect any sensor to the Overo. It is an excellent option for
developing a Brain Computer Interface (BCI) with a direct
connection to the external device e. g. robot or prosthesis
