Biosensors for Health, Environment and Biosecurity
306
deoxycholate (DOCA) was found to be optimal with regard to hemoglobin surface loading,
regeneration and direct reduction of the bound hemoglobin. Unlike their previous work,
blood samples were first incubated with FcBA and then applied on the modified surface.
The boronic acid/diol interaction is much faster in alkaline conditions; on the other hand,
hemoglobin has lower stability at these pHs. Consequently, the optimum pH for incubation
was found to be 8.0. Denaturation of hemoglobin before incubation with FcBA (by heat
treating at 75 °C for 300s) is required for detection of HbA1c and the electrochemical
response of the heme groups and also increases binding with DOCA-modified surface. The
amount of the total hemoglobin bound to the surface is monitored by a quartz crystal
nanobalance (QCN). Upon immobilization of hemoglobin on the electrode surface, the
oscillation frequency of the quartz crystal decreases. The decrease in the frequency is
proportional to the amount of adsorbed total hemoglobin. Fig. 14 shows a typical response
of the QCN upon hemoglobin binding and regeneration of the DOCA-modified
piezosensor. The oscillation frequency decreases after hemoglobin binding, but increases
again after washing loosely bound hemoglobin and returns back to the baseline after
regeneration and removal of bound hemoglobin. More than 30 binding-regeneration cycles
were possible without loss of sensitivity. Fig. 14. Typical QCN response after Hb-binding to the DOCA-modified piezosensor. (A)
Injection of Hb (7.75μM) is followed by (B) washing with buffer (Sörensen phosphate buffer
pH 7.5) and (R) 5 min regeneration using pepsin solution. The dotted line represents the
baseline of the piezoelectric quartz crystal. Before measurement, Hb was incubated at 75 °C
for 300 s (Halámek J. , Wollenberger, Stöcklein, & Scheller, 2007).
These researchers used the same method of square wave voltammetry used in their earlier
work for measurement of the FcBA-bound HbA1c (Fig. 15). To ensure that all HbA1c
molecules are bound to FcBA, they added a 12-fold excess of FcBA to total hemoglobin. Fig.
16 shows the dependence of the current peak height of the SWV on %HbA1c. The standard

approach with that of direct tagging of HbA1c with FcBA described previously shows a 3.6-
fold increase in sensitivity (Fig. 17). Although all the experiments were conducted in a single
day, the standard deviations based on 3 measurements per sample were still high and
accurate detection of HbA1c levels below 5% was still a problem. Fig. 17. Dependence of peak height of the SWV at +300 mV versus Ag/AgCl (1M KCl) on
the HbA1c content in the Hb sample (total Hb 7.75 μM in Sörensen buffer pH 8.0,
preincubated at 75°C). After immobilization of Hb onto the DOCA sensor, either FcBA (○) or
anti-HbA1c Ab and then FcBA (•) was injected. SWV were then measured in stopped flow
(Halámek J. , Wollenberger, Stöcklein, Warsinke, & Scheller, 2007).
Son et al fabricated a disposable biochip for electrochemical HbA1c measurement (Son, Seo,
Choi, & Lee, 2006). They used ferricyanide (K
3
Fe(CN)
6
) as mediator so that the electrons
released from the oxidation of Fe
2+
in hemoglobin were transferred to the electrode by the
ferricyanide/ferrocyanide couple. A schematic view of their %HbA1c measurement
procedure is shown in Fig. 18. The components integrated in the system are a pair of
interdigitated array (IDA) electrodes, HbA1c binding chamber, blood lysis chamber, filter,
micro-pump and microchannel. After plasma separation (1) and red blood cell (RBC) lysis
(2), the total hemoglobin stream branches off into two separate streams: in the lower stream
HbA1c is immobilized on a packed agarose bead containing m-amino-phenylboronic acid
(m-APBA) in the binding chamber and releases hemoglobin, while total hemoglobin flows
in the upper stream (3). The ratio of the resulting electrochemical signals from the lower and
upper streams after passing through the IDA electrodes yields the %HbA1c. Due to the non-
homogeneous distribution of hemoglobin, the instantaneous current varies as a sample


Fig. 20. AFM image the HbA1c/T3BA-SAM immobilized on it (left) along with
corresponding cross-sectional profiles of the spots marked by white circles on the images
(right) (Park, Chang, Nam, & Park, 2008).
Electrochemical determination of selectively immobilized HbA1c on the T3BA SAM is based
on measuring the change in the capability of the gold electrode for electron transfer due to
blocking of the electrode surface by HbA1c after immobilization. This is conducted using
standard HbA1c solutions diluted with a buffered (pH 8.5) solution containing 10 mM 4-
ethylmorpholine in a 3-electrode cell including a gold disk working electrode (0.020 cm
2
),
Ag/AgCl reference electrode and platinum spiral wire counter electrode. The T3BA SAM
has been found to have relatively high electrochemical activity since the charge transfer
resistance R
ct
is small only when it forms on the surface. On the basis of the shape of the EIS
Nyquist plot obtained, the SAM appears to cover the electrode surface uniformly with no
significant defects. The subsequent addition of HbA1c to the system causes the R
ct
value to
increase significantly. As shown in Fig. 21, the ratio of R
ct
obtained in the presence of HbA1c
to that obtained in its absence increases linearly with HbA1c concentration. Similarly, this
ratio varies linearly with %HbA1c in samples with the same total hemoglobin concentration
(Fig. 22). Such linear behaviour makes the T3BA-SAM modified electrode a satisfactory
platform for a HbA1c sensor. On the other hand, these results indicate that the variation of
this signal with HbA1c concentration also depends on total hemoglobin concentration.
Consequently, the total hemoglobin concentration must also be determined to obtain the
HbA1c content. Electrode regeneration can be carried out by washing with a sodium acetate

presence of 0.1 mM ferrocenemethanol (as mediator) and 10 mM glucose (as substrate). The
anodic current measured at +400 mV was chosen as the sensor signal because of stable
current at this potential in the voltammogram. Fig. 24(A) shows voltammograms obtained at
different HbA1c concentrations. As expected, an increase in the HbA1c concentration leads
to a decrease in the resulting current due to less available space for GOx on the electrode.
The corresponding calibration curve for the anodic current at +400 mV as a function of
HbA1c concentration is shown in Fig. 24(B). Although this sensor has the advantage of
signal amplification without the need for pretreatment such as labelling or use of labelled
secondary antibody, incubation of the hemoglobin sample and then GOx solution requires 1
hour and 30 minutes, respectively. In addition, the sensitivity at HbA1c levels below 5% is
not sufficient. Fig. 22. R
ct
ratio obtained at five HbA1c concentrations 20 minutes after sample injection
(Park, Chang, Nam, & Park, 2008).
Qu and coworkers fabricated a micro-potentiometric Hb/HbA1c immunosensor based on
an ion-sensitive field effect transistor (ISFET) using a MEMS fabrication process (Qu, Xia,
Bian, Sun, & Han, 2009). Such ISFET biosensors have numerous advantages such as easy
miniaturization and mass-production and rapid and label-free detection of a wide range of
chemical and biochemical species. The procedure involved modification of the gold working
electrode by electropolymerization of a polypyrrole (PPy)-HAuCl
4
composite followed by
electrochemical synthesis of gold nanoparticles (AuNP) and modification of the gold
reference electrode by applying a PPy film. The presence of AuNP on the surface (confirmed
by FE-SEM) is reported to enhance antibody immobilization. Also, the PPy-AuNp electrode
was electrochemically characterized by cyclic voltammetry and shown to exhibit better
redox reaction reversibility than a PPy electrode. For hemoglobin and HbA1c

Fig. 24. Electrochemical biosensing of HbA1c by using Dend-FPBA electrodes. (A) Cyclic
voltammograms of the backfilling assay between HbA1c and activated GOx at different
HbA1c concentrations in the presence of ferrocenemethanol (0.1mM)in electrolyte with
glucose (10mM)in 0.1MPBS (pH 7.2) at a 5mV/s sweep rate. A voltammogram before
glucose addition is also included for comparison. (B) Calibration curve from the resulting
backfilling assay as a function of target HbA1c concentration. Signal current levels were
masured at +400mV from the background-subtracted voltammograms for respective analyte
concentrations. The mean value from three independent analyses is shown at each
concentration with error bar indicating the standard deviation (Song & Yoon, 2009).
The HbA1c concentration was measured using the same procedure on 10 μL solutions
containing concentrations of 4-18 μg/ml HbA1c in PBS (pH 7.4) Fig. 26 shows a linear dose-
response over this concentration range. Sensor sensitivity and variation coefficient of ΔE
was reported to be 1.5087 mV μg
-1
ml and 24%. The change in response due to the addition

Electrochemical Biosensor for Glycated Hemoglobin (HbA1c)
315
of potential interferents such as immunoglobin G (100 μg/ml), α-fetoprotein (2.5 μg/ml)
and BSA (1%) was found to be less than 9.2%. It was also found that the ΔE of the
hemoglobin sensor decreased about 33.2% after storage at 4°C under dry conditions for 5
days in 100 μg/ml hemoglobin in PBS (pH 7.4). The same trend was observed for a HbA1c
sensor which showed a decrease in ΔE by about 35.1% after storage at 4°C under dry
conditions for 5 days in 8 μg/ml hemoglobin in PBS (pH 7.4). This change in response was
attributed to the slow deactivation of antibodies during storage. Although this sensor has a
short response time (less than 1 min) in comparison to other HbA1c biosensors and low
fabrication costs (in the case of batch produced electrode chips), its low stability and the
relatively high variability of its signal are problems requiring further improvement.
AuNP followed by covalent immobilization on a gold electrode already modified with
mercaptoethylamine-SAM using NHS and EDC. Antibodies were immobilized on the
modified electrode using NHS and EDC as well. SEM images of the modified electrode
showed a more uniform distribution of AuNPs which was attributed to the presence of
SAMs. Electrochemical characterization of the modified electrode using CV and EIS
confirmed that the SAMs had an insulating effect by decreasing the oxidation/reduction
current and increasing the interfacial resistance. Also, the presence of AuNP increased the
electrode sensitivity about 2-fold by raising the surface area-to-volume ratio of the sensor
and making more sites available for antibody immobilization (Fig. 28A).
Measurements of hemoglobin and HbA1c content were conducted on 5 μL samples of
simulated blood solution. Hemoglobin with concentrations of 166.67-570 ng/ml and HbA1c
with concentrations of 1.67-170.5 ng/ml were analyzed. Figs. 28B and C indicate that linear
relations between reagent dose and the electrode response were obtained over the
concentration ranges from 166.67 to 570 ng/ml for hemoglobin and from 50 to 170.5 ng/ml
for HbA1c. Sensor sensitivity was also reported to be 40.42 μV/(ngmL
-1
) and 94.73
μV/(ngmL
-1
) for hemoglobin and HbA1c, respectively. Also, the relative standard deviation
of the measurements (RSD) was 5%. The good linearity of the results was attributed to the

Electrochemical Biosensor for Glycated Hemoglobin (HbA1c)
317
absence of significant interferences from bovine serum albumin, lysis solution, potassium
ions and chloride ions in the simulated blood sample as well as good biocompatibility of the
method and a stable combination with antibodies. In comparison with their previous
sensors based on mixed SAMs, the use of wrapped AuNP arrays increased the sensor
sensitivity from the order of μg/mL to ng/mL and lowered the standard deviation from
above 20% to 5%, while reaching a dilution factor of 150,000 times.

6. References
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RECOGNITION , Vol. 19, pp. 106–180
Berg, A. H., & Sacks, D. B. (2008). Haemoglobin A1c analysis in the management of patients
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987
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Fructosyl Valine. Electroanalysis , Vol. 22, No. 6, pp. 688 – 693
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Halámek, J., Wollenberger, U., Stöcklein, W. F., Warsinke, A., & Scheller, F. W. (2007). Signal
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Sugar Moiety of Immunoglobulin G: Proof of Concept for Glycated Hemoglobin.
Analytical Letters , Vol. 40, pp. 1434–1444
Halámek, J., Wollenberger, U., Stöcklein, W., & Scheller, F. (2007). Development of a
biosensor for glycated hemoglobin. Electrochimica Acta , Vol. 53, pp. 1127–1133

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practice of point-of-care testing (S. 3–12). Lippincott Williams & Wilkins.

Stöllner, D., Stöcklein, W., Scheller, F., & Warsinke, A. (2002). Membrane-immobilized
haptoglobin as affinity matrix for a hemoglobin-A1c immunosensor. Analytica
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A1c Analogon. BIOSENSOR Simposium. TÜBINGEN.
Tsugawa, W., Ishimura, F., Ogawa, K., & Sode, K. (2000). Developement of an Enzyme
Sensor Utilizing a Novel Fructosyl Amine Oxidase from a Marine Yeast.
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Tsugawa, W., Ogawa, K., Ishimura, F., & Sode, K. (2001). Fructosyl Amine Sensing Based on
Prussian Blue Modified Enzyme Electrode. Electrochemistry , Vol 69, No. 12, pp.
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814-825
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Immunosensor for Hemoglobin-A1c Level Detection Based on Mixed SAMs
Wrapped Nano-spheres Array. Biosensors and Bioelectronics , Vol. 26, pp. 2689–2693
14
Electrochemical Biosensors for Virus Detection
Adnane Abdelghani
National Institute of Applied Science and Technology, Charguia Cedex
Tunisia
1. Introduction
The rabies constitutes one of the most dangerous viruses causing many death cases every
year. Each year approximately 55,000 people die of rabies, with high percentage of children
[S et al., 2007; L et al., 2000; FX et al., 1994). High percentages (99%) of the registered cases

N
1
antibodies onto a functionalized gold electrode
with micrometer size. The affinity interaction of the antibody with the specific antigen can

Biosensors for Health, Environment and Biosecurity
322
be measured with a good reproductibility with impedance spectroscopy [M et al., 2008; M et
al., 2008]. The different steps of biosensor conception were characterized by Electrochemical
Impedance Spectroscopy (EIS). The obtained limit detection was better than those obtained
with the others traditional methods for clinical use. The non-specific interaction has been
tested with the Newcastle antigen virus.
2. Experimental set-up
2.1 Specific rabies antibody preparation
Rabies immunoglobulins were produced by horse immunization. The immunization was
carried out using human vaccine “RABIPUR” manufactured by “Chiron Behring Vaccines ″
in Ankleshwar (Gujarat), India. The horses were exposed to a series of injections to increase
vaccine amounts. The immunization period lasted for 105 days (M et al., 2008).
2.2 Specific rabbit antibody (anti-H
7
N
1
) preparation
Three male rabbits were injected sub-cutaneously with different doses of NobilisTM,
INFLUENZA H
7
N
1
vaccine in different periods (15 days, 30 days, 45 days, 65 days). For
each period, quantity of blood were analysed to study the kinetic of the rabbit vaccine

1

2
Z
f
C
π
= (1)
where f is the frequency (in Hz) at which Z is measured.

Electrochemical Biosensors for Virus Detection
323
The complex impedance can be presented as a combination of the real impedance (Z
re
) and
imaginary impedance (Z
im
), Nyquist plot. To fit the measured spectra with the impedance
spectra out of ideal elements, the ideal elements have been replaced with the constant phase
elements (CPE):

K
Z
CPE
α
ω
−
=
(2)


measured spectra of the impedance and phase were analysed in terms of electrical
equivalent circuit model using a Zview modelling programme (Scribrer and associates,
Charlottesville, VA). All electrochemical measurements were carried out at room
temperature and in a faraday cage. More details on electrochemical impedance spectroscopy
can be found in reference (A et al., 2004; M et al., 2008).

Gold electrode

virus antibody
thiol
BSA

Biosensors for Health, Environment and Biosecurity
324
3. Results and discussions
3.1 Avian influenza virus biosensor
First, we study the variation of the impedance spectra (the real part, it means the charge
transfer resistance) of the functionalized gold electrode with different concentration of


Electrochemical Biosensors for Virus Detection
325
R0
Z
CPE
R1

Fig. 3. Electric model

Fig. 4. Impedance spectra of the functionalized gold electrode after the immobilisation of
different antibody concentration

Biosensors for Health, Environment and Biosecurity
326
The interface can be modelized with the electric model shown in figure3. An excellent fitting
between the simulated and experimental spectra was obtained. Fig. 5. Impedance spectra of the functionalized gold electrode before and after addition of
different H
7
N
1
antigen concentration.
The charge transfer resistance increases and reaches a new saturation value that can be
determined with the fitting program. This increase could be attributed to a rearrangement in
the structure of the antibody and a variation of the dielectric constant. The lowest detection
limit that induces a signal variation is equal to 5 μg/ml . This value was lower than the limit
detection obtained with ELISA technique.

We start to characterise the insulating properties of the thiol monolayer on gold surface. The
Nyquist diagram of bare gold (fig.7) presents a half-circle, characteristic of a resistance in
parallel with a capacity and a linear part which appears at low frequencies and which is
assigned to diffusion phenomenon [17]. However, after acid thiol treatment, the diameter of
the half-circle of Nyquist diagram increases clearly and the Warburg impedance was not
observed. This is shows the high insulating properties of the acid thiol (M et al., 2008).
After, impedance measurements were performed after gold surface activation with
EDC/NHS, antibody immobilization and BSA blocking step. After each step, acquisitions
of impedance data in PBS were carried out over 5 decades of frequency. Figure 8 presents
the Nyquist diagram for the various steps of the biosensor development: SAM layer,
antibody layer, blocking with BSA and antigen injection.
We observe that each layer deposition on the gold surface generates an impedance increase.
This increase is due to the change of the electric properties of the gold/electrolyte interface.
3.2.1 Calibration and selectivity
For rabbies detection, Figure 9 presents the calibration of the developed biosensors for
specific and non-specific detection. It shows a dynamic range between 0.1 µg/ml and 4

Biosensors for Health, Environment and Biosecurity
328
µg/ml and a saturation reached at 4 µg/ml. This behaviour can be explained: when the
antigen concentration increases in the electrochemical cell, the number of immobilized
antibodies, but not complexed, decreases and reaches zero when concentration of antigen is
higher than 4 µg/ml. The limit detection of this sensor is about 0.5 µg/ml. This limit
detection is better than the limit detection obtained with the others traditional methods for
clinical use. In order to prove sensor selectivity, the immunosensor was exposed to a
solution containing the Newcastle antigen viruses.

Fig. 7. Nyquist plot for bare gold (small curve) and gold with thiol acid (big curve) in PBS
solution with 1 mM redox couple
As shown in figure 9, there is a little variation of impedance after no-specific antigen

14-Zim[
ΚΩ
.cm
2
]
Zre[ΚΩ.cm
2
]
Bare Gold

-Zim[
ΚΩ
.cm
2
]
Zre[
ΚΩ
.cm
2
]
gold with thiol acid

Electrochemical Biosensors for Virus Detection
329

Fig. 8. Nyquist plot of the different layers of the rabies biosensors


Thiol acide
NHS-EDC
Ac
BSA 0.1%
01234567891011
0
100
200
300
400
500
600
700
800
900
1000
1100
1200
1300
1400

ΔΖ(Ω.
cm
2
)
Antigen concentration (µg/ml)
No specific detection
Specific detection

Biosensors for Health, Environment and Biosecurity

S, Hleli.; A, Abdelghani.; M.A, Maaref

. (2004). Electrical Characterization of a
Thiol SAM on Gold as a First Step of the Fabrication of an Immunosensors based
on a Quartz Crystal Microbalance. Sensors, Vol.4, pp. 104-114.
[13] S, Hleli.; C, Martelet.; A, Abdelghani.; N, Jaffrezic-Renault. (2006). Atrazine analysis
using impedimetric immunosensors based on mixed biotinylated self-assembled
monolayer. Sensors and Actuators B, Vol.113 , pp. 711–717.
[14] A, Abdelghani.; K, Cherif.; M, Maaref. (2006). Impedance Spectroscopy on Xerogel Layer
for Chemical Sensing. Materials Science and Engineering C, Vol. 26, pp. 542 – 545.
[15] A, Tlili.; A, Abdelghani.; S, Ameur.; N, Jaffrezic-Renault. (2006). Impedance
spectroscopy and affinity measurement of specific antibody-antigen interaction.
Materials Science and Engineering C, Vol. 26, pp. 546 – 550.
[16] M, Hnaien.; S, Hleli.; M, Diouani.; I, Hafaid.; W, Hassen.; N, Jaffrezic- Renault.; A,
Abdelghani. (2008). Immobilisation of Specific Antibody on SAM Functionalized
Gold Electrode for Rabies Virus Detection by Electrochemical Impedance
Spectroscopy. Biochemical Engineering Journal, Vol. 39, N0. 3, pp. 443-449.
[17] M, .Diouani.; S, Hleli.; I, Hafaid.; A, Snousi.; A,Ghram.; A,Abdelghani. (2008).
Immobilisation of Specific Antibody on Functionnalised Gold Microelectrode for
Avian Influenza Virus H7N1 Detection. Materials Science and Engineering C, Vol.
28, N0. 5-6 , pp. 580-583.