A Portable Sensing System for Impedance based Detection of

Biotoxin Substances

V. I. Ogurtsov, K. Twomey and J. Pulka

Tyndall National Institute, Lee Maltings, University College Cork, Cork, Ireland

Keywords: Label-free Biosensor, Biotoxin, Immunosensor, Surface Modification, Electrochemical Impedance

Spectroscopy, Portable Instrumentation, Signal Processing.

Abstract: The study describes the development of a portable autonomous biosensing platform for impedance based

detection of biotoxin substances. The platform implements a label-free approach, which is based on detection

of the biosensor interfacial changes due to a bio recognition reaction. The interfacial changes are sensed by

means of Electrochemical Impedance Spectroscopy in a frequency range from 10 Hz to 100 kHz. The platform

comprises of an electrochemical biosensor, portable low-noise mix signal hardware and associated software

incorporating signal processing algorithms for extraction biotarget concentration from the biosensor response.

The biosensor is realized as an on-chip package-free three electrode micro electrochemical cell consisted of

a counter electrode (CE), a reference electrode (RE) and a working electrode (WE) patterned on a single

silicon chip. WE represents an array of 40 um diameter gold disks with 400 m center-to-center distance,

which were undergone of corresponding surface modification for antibody immobilisation. The developed

system was validated by an example of T-2 toxin detection. Performed calibration in the range of 0 – 250 ppm

of T2 toxin concentrations confirmed that the system can provide successful detection of the toxin at the levels

below 25 ppm.

1 INTRODUCTION

Biosensing systems are increasingly used in a range

of different applications including environmental,

clinical, food, agriculture, and security (Bryan et al.,

2013), (Zhang, Du, and Wang, 2015), (Wang, Lu, and

Chen, 2014), (Yong et al., 2015). The conventional

approaches are mainly lab-based and cannot be easily

brought to the point-of-need. Modern microfabricated

biosensors, on the other hand, offer the advantages of

a cost-effective and rapid sample analysis, and

specific and sensitive measurements over the more

traditional methods, which tend to be multi-step (e.g.

ELISA), or to involve sophisticated and expensive

instruments (e.g. HPLC). The ability to apply

semiconductor processing technologies, more

commonly used in the IC industry, in the sensor chip

fabrication (Herzog et al., 2013; Said et al., 2011)

enables large batch production and subsequent

availability of cheap and disposable devices. Typical

microfabricated biosensors incorporate a gold

electrode, or other e.g. platinum, Si

, active layer

upon which different surface chemistries are applied

to form a complete device that is specific and

sensitive to a target of interest. There are different

biosensor technologies Electrochemical, Optical,

Thermal and Piezoelectric. Of these, the

electrochemical biosensors accounted for the largest

share of over 70% of the global biosensors market in

2013, which are expected to maintain their leading

position during the forecast period from 2014 to 2020

(

Transparency Market Research , 2014).

Within the branch of electrochemical sensors,

there are amperometeric, potentiometric,

impedimetric and field-effect transistor (FET)

biosensors (Kafi et al., 2008), (Zhou et al., 2016),

(Ogurtsov, Twomey, and Herzog, 2014). With the

amperometric type biosensors, the changing current

response is monitored with time; for the

potentiometric the voltage is monitored. Both of these

types offer a relatively straight forward measurement

and a rapid analysis method. The impedimetric is

more complex and more sensitive technique. Its

classical implementation is based on the

measurement of an AC current that forms in the

response on the application to the sensor of a

sinusoidal voltage over a set of frequencies.

There is a growing interest in label-free

impedimetric biosensors due to such their advantages

Ogurtsov V., Twomey K. and Pulka J.

A Portable Sensing System for Impedance based Detection of Biotoxin Substances.

DOI: 10.5220/0006169200540062

In Proceedings of the 10th Inter national Joint Conference on Biomedical Engineering Systems and Technologies (BIOSTEC 2017), pages 54-62

ISBN: 978-989-758-216-5

as simplicity of sample preparation, high sensitivity

and the ability to miniaturise the associated

measurement instrumentation (Bryan et al., 2013),

(Rushworth et al., 2014). A further advantage can be

obtained by using of micro-sized electrodes, which

improve the mass transport (and hence the sensor

sensitivity) owing to the occurrence of radial

diffusion over planar diffusion from macro-size

electrodes resulting in an improvement in the signal-

to-noise ratio and reduction in the

iR drop (Arrigan,

2014)

The presented study describes the development

of such portable immunosensing platform capable of

biotoxin detection. The performance of the

biosensing platform was validated by detection of T2-

toxin. It is a trichothecene mycotoxin, which is toxic

to humans and animals. It is naturally occurring

mould byproduct of Fusarium spp. fungus that can be

found in grains such as barley, wheat and oats.

2 EXPERIMENTAL

Description of the Bio Sensing

System

The platform implements a label-free approach,

which is based on detection of the biosensor

interfacial changes due to a bio recognition reaction.

The interfacial changes are sensed by

Electrochemical Impedance Spectroscopy (EIS)

technique (measurement of the biosensor complex

impedance in a wide frequency range) (Lasia, 1999).

The platform structure is shown in Figure 1.

Figure 1: Structure of the biosensing platform.

It comprises of an electrochemical biosensor, low-

noise mix signal portable instrumentation hardware

and associated software incorporating signal

processing algorithms for extraction biotarget

concentration from the biosensor response.

Biosensor Chip

The biosensor is realized as an on-chip package-free

three electrode micro electrochemical cell, which

includes a Pt counter electrode (CE), an Ag/AgCl

reference electrode (RE) and an Au working electrode

(WE) patterned on a single silicon chip as shown in

Figure 2A, Figure 2B and Figure 2C. The WE

represents an array (Figure 2B) consisted of 40 m

diameter recessed gold disks (Figure 2C) arranged in

a hexagonal configuration with 400 m center-to-

center spacing, which were undergone of surface

modification as will be described below. The WE was

placed at the end of the chip to provide its separation

from RE and CE. Such arrangement was

implemented in order to facilitate the bio

functionalization chemistry and prevent both RE and

CE from interference with this process

C B

Figure 2: Photograph of an on-chip package-free three

electrode micro electrochemical cell (A); SEM image of

WE with 40 m microdisk array (B) and SEM image of

recessed gold disk produced by plasma etching of the

silicon nitride isolation layer.

The on-chip micro electrochemical cell was

fabricated at the Central Fabrication Facility at

Tyndall National Institute. The electrodes on the on-

chip microelectrochemical cell were patterned on a Si

substrate of N-Type, <111>-orientation and

fabricated by standard photolithography and lift-off

techniques as described in (Said et al., 2011) and

schematically shown in Figure 3. In brief, firstly, a

silicon oxide layer of 1 μm thickness was thermally

grown on the substrate and alignment marks were

placed on the wafer. Then, the gold working

electrodes, connecting tracks and pads of 150 nm

thick, the platinum counter electrodes of 150 nm

thick and silver reference electrodes of 250 nm thick

were patterned on top of the silicon oxide layer by a

38 mm

10 mm

38 mm

10 mm

A Portable Sensing System for Impedance based Detection of Biotoxin Substances

metal lift-off process. In this process 20 nm of Ti was

used to promote the adhesion of the noble metals to

the wafer that in case of Ag was supplemented by 20

nm of Ni. Then, 500 nm of silicon nitride was

deposited on the whole wafer by plasma-enhanced

chemical vapor deposition. The role of the silicon

nitride is to insulate the connecting tracks from the

solution. Openings for the electrodes and the

connecting pads were obtained by plasma etching.

Silver–silver chloride reference electrodes were

prepared with help of chemical oxidation by

immersion of the wafer in a 10 mM FeCl

solution for

50 s. The metal was then lifted and a resist layer was

spin coated on top of the wafer to protect the

electrodes during the dicing of the wafer into

individual chips.

Figure 3: Fabrication process flow for the three electrode

on-chip micro electrochemical cell.

Method and Chemicals

All electrochemical experiments were performed in a

Faraday cage, and each measurement used for

obtaining sensor calibration was carried out three

times. EIS measurements were performed by the

described system over a frequency range from 10 Hz

to 100 kHz at an applied potential of DC bias of 0.2

V and an AC amplitude of 10 mV. The frequency

range was defined by hardware specification and

application objective to determine analyte

concentration. The appropriate bias potential was

determined from cyclic voltammetry over a range 0V

to 0.6 V at a scan rate of 100 mV/s. All chemicals

used in this work were purchased from Sigma Aldrich

Ireland Ltd. and utilised as received.

Surface Modification Procedure

The working electrode of the microchemical cell was

undergone of surface modification for attachment of

antibodies capable of receipting the selected biotarget

according to a scheme presented in

Figure 4

Figure 4: Surface modification scheme.

This procedure consisted of the following steps:

Electrode Surface Pretreatment. The fabricated

chips were first plasma-cleaned for 10 minutes and

then immersed in HCl:MeOH (1:1, v/v) solution for

15 minutes. They were then sonicated in acetone and

isopropyl alcohol for 5 minutes each; rinsed with

copious amounts of DI water and dry under a stream

of nitrogen.

Silanisation of the WE surfaces was carried out in

3% APTES in MeOH:DI water (19:1) solution for 30

minutes at room temperature. The electrodes were

then rinsed sequentially with MeOH and DI water

before left cure in the oven (dust free) for 15 minutes

at 120°C.

Surface Activation (Cross-Linker Attachment).

Following the curing steps, the silanised electrodes

were immediately immersed in 18 mL of DMF

solution containing 2 mL of 10% pyridine and 0.098

g 1,4-phenylene diisothiocyanate (-PDITC)

(produces 25mM PDITC) for 2 hours. The electrodes

BIODEVICES 2017 - 10th International Conference on Biomedical Electronics and Devices

were washed sequentially with DMF and DCE and

dry under N

Antibody Immobilisation. For this modification,

anti-T2 toxin antibody was diluted in 0.1 M sodium

borate pH 9.3. The working electrodes were

immersed in 100 μL of diluted antibody solution for

2 hours at RT, wrapped in aluminium foil. Then the

electrodes were removed, rinsed with DI water and

dried with N

Instrumentation

The instrumentation block diagram is shown in

Figure

. It represents a two channels EIS system where each

channel contains a signal generator providing in-

phase and quadrature AC signals, transimpedance

amplifier and two (in-phase and quadrature) detection

channels, each consisted of a mixer and a low-pass

filter. These two paths are a sensor channel for the

actual measurement and a reference channel to

compensate for the background signal. The sensor

connected to the reference channel should be not

specific to the target analyte thereby capable of

providing information on the sensor background

signal resulting from varying parameters such as

temperature, pH, nonspecific binding etc. If the

background signal is of negligible value, the

reference channel can be used as the second sensing

channel for another bio target.

A known AC voltage that is generated by the

signal generator is applied across the sensor and the

sensor current is converted to a voltage by the

corresponding transimpedance amplifier. The voltage

is applied across the sensor and the sensor is

amplified and demodulated in the I/Q demodulator

yielding real and imaginary impedance components.

To increase the dynamic range of the module, 2-

bit electronically controlled instrumentation

amplifiers are included in the channel structure. To

facilitate fine module adjustment and calibration,

level shifters on the base of I

C controlled DAC are

introduced into the circuit.

The biosensing system hardware is realized

consisted of two modules: Potentiostat and Front-End

Amplifiers (PFEA) and Signal Processing and

Microcontroller (SPM) units.

The PFEA (Figure 5) is designed as a separate

small size unit in order to locate it close to a reservoir

where the biosensor tests the sample solution.

Figure 5: Potentiostat and Front-End Amplifiers module:

photographs of the packaged unit with the sensor chip (A)

and the unit PCB (B).

Figure 6: Instrumentation block diagram.

A Portable Sensing System for Impedance based Detection of Biotoxin Substances

This shortens the length of the connection wires,

which link the biosensor and front-end amplifier,

therefore decreases the leakage current and attenuates

the noise and electromagnetic interference. The

PFEA unit consists of a potentiostat, a temperature

sensor and two transimpedance amplifiers providing

current to voltage transformation of the sensor and

reference (calibration) signals. The potentiosat was

realized on three operational amplifiers OPA2211.

Two identical transimpedance amplifiers were

assembled from OPAMP ADA4647 and

instrumentation amplifier AD8253. The unit was

implemented on a small four-layer PCB of 25 mm x

35 mm dimensions.

The Signal Processing and Microcontroller

(SPM) unit (Figure 7) includes mixing and Analog

Signal Processing (ASP), microcontroller and

peripheral (C) and Power Supply (PS) blocks.

Figure 7: Photograph of Signal Processing and Micro-

controller unit.

The ASP block consists of Input, Switch, Level-

Shifter and four identical Mixer/Low Pass Filter

(LPF) lock-in circuits, which form two identical

analog processing channels. Four-channel 16-bit

digital to analog converters (DAC) incorporated into

the SPM circuit is used to adjust DC levels in the SPM

unit. The block was assembled from AD630 (Mixer),

UAF42 (Low Pass Filter), ADG1419 (Switch),

DAC8574 (DAC with buffered voltage output and

C compatible two wire serial interface), AD8253

(Instrumental amplifier with 2-bit electronically

controlled gain) and LM4140 (1.024 V voltage

reference).

The C block is based on a high performance 8-

bit ATxmega128A1U microcontroller. It realises

functions of instrumentation control, data acquisition

and communication. It also contains the signal

generator and four 16-bit DACs. The signal generator

is based on AD9854 chip, which is a

DDS synthesizer

capable of parallel generation of precision sine and

cosine signals.

The PS block supplies stabilised low noise

voltages of +8V, -8V, +3.3V and -3.3V to power the

signal generators, microcontroller, ASP and PFEA

system units. The PS block can be powered from a

voltage source in the range of +5V – +12V.

These three blocks formed the main body of the

instrumentation were implemented in three separate 4

layer FR4 PCBs of 51 mm x 90 mm dimensions,

which were integrated in a stack manner taking 55

mm of height as shown in Figure 7.

Signal Processing and Corresponding

Software

The associated with instrumentation software

consists of two highly interrelated parts written for a

host computer and an ATxmega128A1U Atmel

microcontroller. The software ensures

communication between the instrumentation and the

host computer, control of the instrumentation settings

and its operation, implementation of the dedicated

procedure of biosensor impedance spectra

measurement and their signal processing. A smoothed

differential impedance spectrum is used for extraction

an analytical signal that is applied for biosensor

calibration and quantification of the target analyte

concentration. A developed automated procedure for

analytical signal extraction includes the following

steps: smoothing of the initial biosensor impedance

spectrum by Kernel smoothing procedure;

subtraction of a reference background spectrum from

the measured spectra; and finally finding the

maximum of the imaginary component of the

differential impedance spectrum, which is taken as

the analytical signal to be used for biosensor

calibration and calculation of the target concentration

(based on the biosensor calibration and extracted

value of the analytical signal).

The PC software was implemented as user-

friendly multi-tab GUI software where each tab is

associated with the program window contained

means for solutions of assigned tasks. There are three

tabs/windows, namely ‘Protocol’, ‘System

configuration’, and ‘EIS’.

The tab/window ‘Protocol’ allowed for careful

design of frequency sweep measurement protocol for

each frequency point by applying values of frequency

and measurement number with corresponding

adjustment of all electronically controlled hardware

parameters.

BIODEVICES 2017 - 10th International Conference on Biomedical Electronics and Devices

Figure 8: Screen short of System configuration PC software tab/window.

The tab/window ‘System configuration’ shown in

Figure 8 provides communication control including

received and transmitted data stream visualisation,

detection and control of communication and stream

errors, data extraction from communication stream,

instrumentation and measurement setting control and

configuration and system operability testing. Tools

for system operability testing include Self-Test

command, which starts Self-test measurement

procedure followed by reporting about the

measurement results in the ‘Test parameters’ plot and

in the ‘Test parameters’ table in ‘Hardware

Configuration’ panel. The program monitors these

parameters before each impedance spectrum

measurements that allows for easy identification of

the hardware and the biosensor operability problem.

3 RESULTS AND DISCUSSION

The tab/window ‘EIS’ is the main instrumentation

software part that is designed to provide all raw

impedance measurements and signal processing

results in different forms as shown in Figure 9. These

include raw sensor impedance measurements

presented by Nyquist (Imaginary impedance part vs.

Real impedance part - ‘Zim vs Zre’) and Bode

(Imaginary impedance part vs. Frequency - ‘Zim vs

F’, Real impedance part vs. Frequency - ‘Zre vs F’

and Phase vs. Frequency - ‘Phase vs F’) plots as well

as the processed Nyquist spectra obtained after

Kernel smoothing and subtraction of the reference

background spectrum from the measured spectra with

analytical signal extraction (see the plot ‘Zim vs Zre

processed’ where the analytical signals extracted are

highlighted by red points on the corresponding

processed impedance spectra) and calculated analyte

concentration in respect to the biosensor calibration

(plot ‘Anal. Signal vs. Concentration’) in both

graphical and table forms.

The experimental results presented in Figure 9

were achieved with T-2 toxin biosensor described

above, which was used for the developed system

validation. Initial and differential raw and smoothed

impedance spectra were obtained for five T2-toxin

concentrations 0, 25, 50, 100 and 250 ppm. In order

A Portable Sensing System for Impedance based Detection of Biotoxin Substances

Figure 9: Screen short of EIS PC software tab/window.

to provide better distinguishability of the spectra the

measurements were made without repetitions thus

zero measurement errors are shown in the

corresponding table below the “Electrochemical

measurement” panel. As one can see, the initial

impedance spectra were subjected to noise and

application of Kernel’s smoothing allowed for

effective noise suppression. Due to a relatively high

low frequency boundary of the sweep range (10 Hz)

the obtained impedance spectra accounted only for

high and middle frequency parts of a depressed

semicircle response of a microdisc electrode array

associated with semi-infinite radial spherical

diffusion. When toxin concentration increases, real

and imaginary impedance parts both also increase that

reflects a growth of a layer with targeted species

captured by antibodies due the antibody-antigen

binding reaction. If toxin concentration exceeds the

certain level this impedance increase is saturated due

to a depletion of the free antibodies capable of

biorecognition reaction.

As follows from Bode plots the most difference

between spectra corresponding to the different toxin

concentrations is located in a low frequency range

below 250 Hz. Here, the imaginary impedance values

are about one order less than their real counterparts;

they are also more sensitive to variation of T2 toxin

concentration. If frequency increases, both

impedance components decrease but the imaginary

impedance stays relatively constant at the frequencies

below 250 Hz where the influence of the toxin

concentration on the biosensor impedance is the most

noticeable. This frequency behaviour of the

imaginary impedance of the biosensor leads to the

fact that frequency dependence of the impedance

phase represents unimodal smooth function with

maximal values in the range of 56 – 62 degrees

around the frequency of 1 kHz.

The dependence of imaginary impedance against

toxin concentration becomes more apparent if to

study the differential impedance spectra obtained by

subtraction of the spectrum at zero concentration,

BIODEVICES 2017 - 10th International Conference on Biomedical Electronics and Devices

which is used as the reference background spectrum,

from the impedance spectra corresponded to nonzero

toxin concentrations. These smoothed differential

impedance spectra are shown in the plot ‘Zim vs Zre

processed’. There are also highlighted by red points

the maximum values of the imaginary impedances on

the corresponding processed impedance spectra.

These values were used as the analytical signals

related to calibration curve of the developed label-

free impedance immunosensor. This curve is shown

in the plot ‘Anal. Signal vs. Concentration’ together

with the analytical signals automatically extracted

from the smoothed differential impedance spectra

presented by the red squares. The calibration curve

represents a nonlinear function that reflects the

competitive nature of the antibody-antigen binding

reaction. Parameters of this calibration function

defined by application of logarithmic regression

analysis are given in Table 1.

Table 1: Calibration parameters of the biosensing system

with T2 toxin biosensor.

Function Slope R

Error

y= a·ln(x+b)+c a= 2.412·10

0.977 7.144·10

Based on these parameters and the extracted values of

analytical signals corresponding T2 concentrations

can be determined as presented in the right table in

the Table panel in Figure 9. As follows from the

calibration curve the developed biosensing portable

system can successfully detect biotoxin at the levels

below 25 ppm.

4 CONCLUSIONS

A portable biosensing platform for impedance based

detection and quantification of biotoxin substances

capable of operation in autonomous mode has been

developed. The platform implements a label-free

approach, which is based on detection of the

biosensor interfacial changes due to a bio recognition

reaction. The interfacial changes are sensed by means

of Electrochemical Impedance Spectroscopy in a

frequency range from 10 Hz to 100 kHz. The platform

comprises of an electrochemical biosensor, portable

low-noise mix signal hardware with embedded

microcontroller and associated software

incorporating hardware and measurement control

together with signal processing algorithms for

extraction biotarget concentration from the biosensor

response. The biosensor is realized as an on-chip

package-free three electrode micro electrochemical

cell consisted of a Pt counter electrode, an Ag/AgCl

reference electrode and an Au working electrode

patterned on a single silicon chip. The WE represents

an array of 40 m diameter gold disks with 400 m

center-to-center distance, which were undergone of

surface modification for antibody immobilisation.

The surface modification was based on APTES –

PDITC procedure. The instrumentation was realised

as a two channel EIS system, which used in-phase and

quadrature demodulation of AC signal for extraction

of the real and imaginary impedance components.

The biosensing system hardware consisted of two

modules: Potentiostat and Front-End Amplifiers

(PFEA) and Signal Processing and Microcontroller

(SPM) units. The corresponding software capable of

autonomous operation contains two interrelated parts

written for the host computer and the embedded

microcontroller. An automated procedure for

analytical signal extraction consists of: smoothing of

the initial biosensor impedance spectrum by Kernel

smoothing procedure; subtraction of a reference

background spectrum from the measured impedance

spectrum; and extraction from differential

impedance spectrum the maximum of the imaginary

component, which is used as an analytical signal for

biosensor calibration and calculation of the target

analyte concentration. The PC software was

implemented as user-friendly multi-tab GUI software

where each tab is associated with the program

window. These are ‘System configuration’,

‘Protocol’ and ‘EIS’ tabs/windows. They provide

communication, instrumentation and measurement

setting control, and system operability testing; allow

for design of impedance measurement protocol for

each frequency sweep point; and present results of

EIS and analyte concentration quantification in

graphical and table forms. The developed biosensing

system was validated with the developed on-chip T2

toxin biosensor. Calibration of the system was

obtained for five toxin concentrations from 0 ppm to

250 ppm. It showed that the developed portable

biosensing system can provide successful detection of

the biotoxin at the levels below 25 ppm.

ACKNOWLEDGEMENTS

Financial support of this work by European

Commission projects FP7-SEC-2011.3.4-2

“HANDHOLD: HANDHeld OLfactory Detector”

and H2020-NMP-29-2015 “HISENTS: High level

Integrated SEnsor for NanoToxicity Screening” is

gratefully acknowledged.

A Portable Sensing System for Impedance based Detection of Biotoxin Substances

REFERENCES

Bryan, T., Luo, X., Bueno, P.R. and Davis, J.J. (2013) ‘An

optimised electrochemical biosensor for the label-free

detection of C-reactive protein in blood’, Biosensors

and Bioelectronics, 39(1), pp. 94–98.

Zhang, W., Du, Y. and Wang, M.L. (2015) ‘Noninvasive

glucose monitoring using saliva nano-biosensor’,

Sensing and Bio-Sensing Research, 4, pp. 23–29.

Wang, X., Lu, X. and Chen, J. (2014) ‘Development of

biosensor technologies for analysis of environmental

contaminants’, Trends in Environmental Analytical

Chemistry, 2, pp. 25–32.

Yong, D., Liu, C., Zhu, C., Yu, D., Liu, L., Zhai, J. and

Dong, S. (2015) ‘Detecting total toxicity in water using

a mediated biosensor system with flow injection’,

Chemosphere, 139, pp. 109–116.

Herzog, G., Moujahid, W., Twomey, K., Lyons, C. and

Ogurtsov, V.I. (2013) ‘On-chip electrochemical

microsystems for measurements of copper and

conductivity in artificial seawater’, Talanta, 116, pp.

26–32.

Said, N.A.M., Twomey, K., Ogurtsov, V.I., Arrigan,

D.W.M. and Herzog, G. (2011) ‘Fabrication and

Electrochemical characterization of micro- and

Nanoelectrode arrays for sensor applications’, Journal

of Physics: Conference Series, 307, p. 012052.

Transparency Market Research (2014) 'Global Biosensors

market - industry analysis, size, share, growth, trends

and forecast 2014-2020', Available at:

http://www.transparencymarketresearch.com/biosenso

rs-market.html (Accessed: 15 December 2016).

Kafi, A.K.M., Lee, D.-Y., Park, S.-H. and Kwon, Y.-S.

(2008) ‘Potential application of hemoglobin as an

alternative to peroxidase in a phenol biosensor’, Thin

Solid Films, 516(9), pp. 2816–2821.

Cheng, S., Hideshima, S., Kuroiwa, S., Nakanishi, T. and

Osaka, T. (2015) ‘Label-free detection of tumor

markers using field effect transistor (FET)-based

biosensors for lung cancer diagnosis’, Sensors and

Actuators B: Chemical, 212, pp. 329–334.

Ramesh, R., Puhazhendi, P., Kumar, J., Gowthaman, M.K.,

D’Souza, S.F. and Kamini, N.R. (2015) ‘Potentiometric

biosensor for determination of urea in milk using

immobilized Arthrobacter creatinolyticus urease’,

Materials Science and Engineering: C, 49, pp. 786–

792.

Zhou, Y., Tang, L., Zeng, G., Zhang, C., Xie, X., Liu, Y.,

Wang, J., Tang, J., Zhang, Y. and Deng, Y. (2016)

‘Label free detection of lead using impedimetric sensor

based on ordered mesoporous carbon–gold

nanoparticles and DNAzyme catalytic beacons’,

Talanta, 146, pp. 641–647.

Ogurtsov, V.I., Twomey, K. and Herzog, G. (2014)

‘Development of an Electrochemical Sensing System

for Environmental Monitoring of Port Water Quality to

Integrate On-board an Autonomous Robotic Fish.’, in

Hashmi, S. (ed.) Comprehensive materials processing.

Volume 13: Sensor Materials, Technologies and

Applications. Oxford, United Kingdom: Elsevier

Science, pp. 317–351.

Bryan, T., Luo, X., Bueno, P.R. and Davis, J.J. (2013) ‘An

optimised electrochemical biosensor for the label-free

detection of C-reactive protein in blood’, Biosensors

and Bioelectronics, 39(1), pp. 94–98.

Rushworth, J.V., Ahmed, A., Griffiths, H.H., Pollock,

N.M., Hooper, N.M. and Millner, P.A. (2014) ‘A label-

free electrical impedimetric biosensor for the specific

detection of Alzheimer’s amyloid-beta oligomers’,

Biosensors and Bioelectronics, 56, pp. 83–90.

Arrigan, D.W.M. (2004) ‘Nanoelectrodes, nanoelectrode

arrays and their applications’, The Analyst, 129(12), pp.

1157-1165.

Lasia, A. (1999) ‘Electrochemical Impedance Spectroscopy

and Its Applications’, in Conway, B.E., Bockris, J., and

White, R.E. (eds.) Modern Aspects of Electrochemistry,

Volume 32. New York: Kluwer Academic/Plenum

Publishers, pp. 143–248.

BIODEVICES 2017 - 10th International Conference on Biomedical Electronics and Devices