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Sensors and Actuators B 134 (2008) 95–103
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Sensors and Actuators B: Chemical
journal homepage: www.elsevier.com/locate/snb
Conductimetric immunosensor for atrazine detection based on antibodies
labelled with gold nanoparticles
Enrique Valera a , Javier Ramón-Azcón b , F.-J. Sanchez b , M.-P. Marco b , Ángel Rodrı́guez a,∗
a
Micro and Nano Technologies Group (MNTg), Departament d’Enginyeria Electrònica, Universitat Politècnica de Catalunya, C/. Jordi Girona 1-3 Campus Nord,
Mòdul C4, Barcelona 08034, Spain
b
Applied Molecular Receptors Group (AMRg), IIQAB-CSIC, CIBER of Bioengineering, Biomaterials and Nanomedicine, Jordi Girona 18-26, 08034 Barcelona, Spain
a r t i c l e
i n f o
Article history:
Received 22 February 2008
Received in revised form 12 April 2008
Accepted 14 April 2008
Available online 24 April 2008
Keywords:
Immunosensor
Gold nanoparticles
Antibodies
Interdigitated ␮-electrodes
Conductive measurements
Impedance spectroscopy
Atrazine
Wine matrix effect
Food safety
a b s t r a c t
A novel conductimetric immunosensor for atrazine detection has been designed and developed.
This immunosensor is mainly based on antibodies labelled with gold nanoparticles. Additionally, the
immunosensor consists of an array of two coplanar non-passivated interdigitated metallic ␮-electrodes
(ID␮E) and immunoreagents specifically developed to detect this pesticide. The chemical recognition layer
was covalent immobilized on the interdigital space. Immunochemical detection of the concentration of
atrazine is achieved by a competitive reaction that occurs before the inclusion of the labelled antibodies.
It is shown that the gold nanoparticles provide an amplification of the conductive signal and hence makes
possible to detect atrazine by means of simple DC measurements.
The conductimetric immunosensor and its biofunctionalization steps have been characterized by chemical affinity methods and impedance spectroscopy.
This work describes the immunosensor structure, fabrication, physico-chemical and analytical characterization, and the immunosensor response using conductivity measurements. The immunosensor developed detects atrazine with limits of detection in the order of 0.1–1 ␮g L−1 , far below the maximum residue
level (MRL) (100 ␮g L−1 ) established by European Union (EU) for residues of this herbicide in the wine.
Although in this paper the competitive reaction occurs in buffer, an initial study of the wine matrix
effect is also described.
© 2008 Elsevier B.V. All rights reserved.
1. Introduction
In the recent years, modern chemical analysis has been revolutionized by the electrochemical biosensors because of their
accuracy, easy use, high efficiency, possibility of portability and
miniaturisation, and because they offer fast (few seconds) response
times, allow a rapid and permanent control and a direct transduction of the biomolecular recognition event into electronic signals
[1–7]. The ability to detect very small amounts of the target substance, the use for continuous monitoring, mass fabrication, low
cost, and decentralized infield analysis are other important features
of these electrochemical sensors.
It is well known that binding of biomolecules to the surface
of solid supports produce changes in the electrical properties in
the vicinity of the electrodes [8]. However, an important disadvantage of these electrochemical sensors is that the impedance
changes due to biomolecular recognition are generally very small.
∗ Corresponding author. Tel.: +34 934016876.
E-mail address: [email protected] (Á. Rodrı́guez).
0925-4005/$ – see front matter © 2008 Elsevier B.V. All rights reserved.
doi:10.1016/j.snb.2008.04.023
Sensors based on interdigitated ␮-electrodes (ID␮Es), have recently
demonstrated important improvements in sensitivity of electrochemical detection [9–11], and are also used in this work.
Except for few transducing principles already well established (i.e., surface plasmon resonance, SPR) [12,13], most of the
immunosensors reported until now, rely on the use of labels to
reach the necessary detection limits required by the legislation.
Typically, enzymes, radioactive isotopes, fluorescein, metal or
semiconductor nanoparticles, or inorganic metal catalysts conjugated to biorecognition events are used as amplifying labels
[14,15]. Among the most recently used labels, for the determination
of biomolecules and some biological metabolites, gold nanoparticles can be found [4,15–22]. The unique properties at nanoscale
dimension of these particles have attracted widespread attention
in their utilization for the bioassay, especially for electrochemical
detection, where the gold particles can be used as conductive
pathway for electron transfer, improving the electrochemical
reactions at a low potential [23–26]. Gold nanoparticles have been
extensively and successfully applied for the immobilization and
study of various kind of biomolecules and macromolecules, such
as DNA, enzymes, other proteins and antibodies [27–29], as well
as for enhancing the binding signals and improving the sensitivity
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E. Valera et al. / Sensors and Actuators B 134 (2008) 95–103
[17]. Gold nanoparticle-labelled technique offers a large number
of advantages such as easy preparation, very large surface area,
excellent biocompatibility, simplicity, accuracy and non-pollution
[18,20,27,30].
Although in this paper, gold labels are used to increase the conductive signal and improve the detection sensitivity (gold labels
enhance the signal of antibodies and antigens immobilized on the
electrode surface), our immunosensor does not require any kind
of amplification to reach the detection limits required by the legislation in line with our previously reported results [31–33]. Our
main objective in this work is to explore whether atrazine can be
detected by means of simple and inexpensive DC measurements
provided gold nanoparticles are included as labels in the biosensor.
In this work we report the development of an immunosensor
based on antibodies labelled with gold particles (40 nm), where
conductive measurements are applied to detect atrazine. The antibodies labelled with gold particles are named secondary antibodies
because they are included in the system after the competitive reaction has been completed as described in detail in Refs. [31–33].
The method described below does not use any redox mediator and relies on the direct detection of the immunochemical
competitive reaction between the pesticide and a haptenizedprotein immobilized on interdigitated ␮-electrodes for a specific
antibody.
2. Materials and methods
2.1. Instrumentation
Electrochemical and conductive measurements were carried out
at room temperature in a probe station (Faraday cage) KARL SUSS.
Impedance analyses were performed using an Agilent 4294A Precision Impedance Analyzer and conductive measurements were
performed using an Agilent 4156C Semiconductor Parameter Analyzer.
The pH and the conductivity of all buffers and solutions, the
absorbances, and the competitive curves were measured, read and
analyzed by means of the same apparatus and techniques used in
the previous works [31–33].
Data shown correspond to the average between 5 and 10 replicates per concentration of atrazine. The reproducibility of the
measurements was checked in all experiments shown in this paper,
measuring each device at least two times in different days.
2.2. Chemicals and immunochemicals
All the chemicals and immunochemicals used before the antibody labelled with gold nanoparticles is deposited, are the same
described in detail in Ref. [31]. Antibody anti-IgG labelled with gold
nanoparticles of 40 nm was prepared following a standard protocol
[34] obtaining a solution with an optical density of 3.0 A520 .
2.3. Buffers
Unless otherwise indicated the buffers used are the same
described in detail in Ref. [33].
2.4. Experimental methods
2.4.1. Fabrication and pre-treatment of the arrays of
interdigitated -electrodes
Thin Au/Cr (∼200-nm thickness) interdigitated ␮-electrodes
(ID␮Es) with 10 ␮m pitch were patterned on a Pyrex 7740 glass
substrate (purchased from Präzisions Glas & Optik GmbH, 0.7 mm
(±0.05) thickness) and finally separated in small arrays (0.99 cm2
area) of six devices. Every interdigitated device had two electrodes.
Metal deposition was performed by sputtering and the ID␮Es were
patterned by a photolithographic metal etching process.
Before functionalization, the surfaces of the samples were first
cleaned as described also in Ref. [33].
2.4.2. Immunosensor surface activation
After the pre-treatment explained above, surface activation took
place in two steps to modify selectively the gold electrodes and the
Pyrex substrate. The activation performed is the same detailed in
Ref. [33].
2.4.3. Antigen immobilization
The covalent immobilization of the competitor antigen was performed on the interdigitated ␮-electrodes surface via the side chain
amino groups of lysines or arginines with the epoxy groups on the
device surface using the same procedure explained in Ref. [33].
2.4.4. Assay development
The electrodes arrays were immersed (500 ␮L/array) in the
atrazine standards prepared in PBST (0.32–2000 ␮g L−1 ), followed
by the antisera Ab 11 (0.25 ␮g mL−1 in PBST, 500 ␮L/array). After
30 min of incubation time at RT, the array was washed as in Ref. [33].
A solution of anti-IgG-gold (1/1000 in PBST) was added (1 mL/array)
and incubated for 30 min more at RT. The arrays were washed again,
and immersed in PBS (1.6 ␮S cm−1 , 1 mL/array).
2.4.5. Protocol of the impedance measurements
Impedance measurements were carried out at room temperature and performed for every functionalization step, in order to
characterize the system modifications. No redox mediator was
used in the devices presented in this work. The two electrodes
were covered by a diluted PBS solution with a conductivity of
1.6 ␮S cm−1 and connected to the input of an Agilent 4294A
Precision Impedance Analyzer by means of standard probe tips.
Measurements were taken in the 40 Hz to 1 MHz frequency range
using 0 V of polarization potential and a modulation voltage of
25 mV amplitude. All impedance measurements were performed
in a Faraday cage.
2.4.6. Protocol of the conductance measurements
As for impedance measurements, the conductance measurements were also carried out at room temperature without any
redox mediator and in a Faraday cage. The two electrodes were covered by a diluted PBS solution with a conductivity of 1.6 ␮S cm−1
and connected to the input of an Agilent 4156C Semiconductor
Parameter Analyzer by means of standard probe tips. Conductivity was measured to +100 mV, from +90 to +110 mV sweep bias and
to +25 mV, from +22.5 to +27.5 mV sweep bias. These conductive
measurements were performed after the incubation step of the secondary antibody labelled with gold. The standard curves obtained
from these conductive measurements, were fitted to the following
four-parameter equation [35]:
Y=
A−B
1 + (x/C)D
+ B,
where A is the maximal absorbance, B is the minimum absorbance,
C is the concentration producing 50% of the maximal absorbance,
and D is the slope at the inflection point of the sigmoid curve.
3. Results and discussion
In this work ID␮Es were chosen because they provide advantages which were described in Ref. [31], and in addition, the use
E. Valera et al. / Sensors and Actuators B 134 (2008) 95–103
97
Fig. 1. Interdigitated ␮-electrode array (six devices) fabricated for the conductive immunosensor (a) optical image and (b) SEM image.
of electrodes arrays offers the possibility to obtain several replicates per assay. Fig. 1 shows an image of one of the interdigitated
␮-electrodes arrays fabricated.
Since our ID␮Es are metallic (and not protected by an insulator),
the PBS conductivity had to be adjusted (1.6 ␮S cm−1 ) by dilution
to provide accurate and sensitive impedance measurements.
The interdigitated ␮-electrodes were immersed into a solution
of N-acetylcysteamine, in order to protect the fingers for nonspecific adsorption. Later, the ID␮Es were immersed into a solution
of 3-glycidoxypropyl trimethoxysilane (GPTS), in order to functionalize their surface.
In this conductimetric immunosensor, the detection of a small
number of molecules is performed under competitive conditions
involving the competition between the free antigen (analyte) and
a fixed amount of coated antigen for a limited amount (low concentration) of a specific primary antibody (Ab1 ). At the end of the
reaction, the amount of specific Ab1 captured on the ID␮E surface (inter-digits space) and hence the free antigen (analyte) can
be determined. Finally, the conductive signal is amplified by means
of a secondary antibody labelled with gold nanoparticles (Ab2 ).
3.1. Immunosensor impedance characterization
The immunosensor functionalization steps have been analyzed
by means of impedance spectroscopy (IS). The physical interpretation of the distributed elements in an equivalent circuit is essential
in understanding and interpreting of impedance spectra.
3.1.1. Equivalent circuit to model impedance data
The measurement of the impedance as a function of the frequency is better understood using an electronic equivalent circuit
which is a based on the general electronic equivalent circuit of
Randles and Ershler [36–38] modified [31–33].
In summary the equivalent circuit includes six elements:
(i)
(ii)
(iii)
(iv)
(v)
(vi)
contact resistance, Rc ;
capacitance of the ID␮E, CID␮E ;
ohmic resistance of the solution, Rs ;
polarization resistance, Rp ;
Warburg impedance from the diffusion, Zw ;
double-layer capacitance, Cdl .
This equivalent circuit correspond to the impedimetric
immunosensor studied previously [31–33]. Strictly, the equivalent
circuit for the conductimetric immunosensor should be modified
to take into account the DC conductance, because the Warburg
impedance and double-layer capacitance are connected in series
with the solution resistor and since both of them have an infinite
module of impedance (in DC), the DC circuit for this equivalent
circuit would be zero. But nevertheless since the contribution of
gold nanoparticles to the increase of conductance is neglectible in
the margin of frequencies from 40 Hz to 1 MHz the equivalent circuit corresponding to the impedimetric immunosensor has been
approximated to the conductimetric immunosensor, only for the
fitting values in this range of frequency.
3.2. Characterization of immunosensor response
The chemical changes on the immunosensor surface happen in
five steps, two initial steps for the immunosensor surface functionalization and three final steps for the immunosensor reaction
itself:
(i) Step I: protection of interdigitated ␮-electrodes with Nacetylcysteamine;
(ii) Step II: immunosensor surface functionalization with GPTS;
(iii) Step III: covalent immobilization of the antigen on the ID␮E;
(iv) Step IV: specific primary antibody (Ab1 ) capture in the competition step;
(v) Step V: secondary labelled with gold antibody (Ab2 ) capture.
These steps are schematically shown in Fig. 2. The gold surface
modification was accomplished using the thiol-chemistry. Likewise, N-acetylcysteamine was used in order to cover the gold
electrodes and to protect the immunosensor from undesired nonspecific absorptions. The resulting Au–S bond granted the stability
of the deposited surface layer. In this case, the surface texture
of the ID␮E defined the template for deposition of layers, since
the gold fingers have been deposited on a solid support such as
glass with the necessary controlled geometry. For the case of the
glass material that serves as support, the most used activation
procedures are based on the silane-chemistry. Thus, on a second
step the Pyrex substrate was derivatized with 3-glycidoxypropyl
trimethoxysilane. The epoxy group provided the necessary reactivity for further attachment of the immunoreagents through a
nucleophilic attack of the amino groups of the lysine residues of
the antigen.
Several different known values of atrazine concentration were
used in the competition step. The sensor implemented in this work
will be sensitive to the chemical changes produced at the sensor
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E. Valera et al. / Sensors and Actuators B 134 (2008) 95–103
Fig. 2. Schematic diagram of the complete assay system performed on the ID␮Es: Step I, protection of interdigitated ␮-electrodes with N-acetylcysteamine; Step II,
immunosensor surface functionalization with GPTS; Step III, covalent immobilization of the antigen on the ID␮E; Step IV, specific primary antibody (Ab1 ) capture in the
competition step; Step V, secondary labelled with gold antibody (Ab2 ) capture. In the Step IV, an amount of the specific antibody (Ab1 ) is bounded on the coated antigen layer,
whereas other amount is evacuated of the ID␮Es, this amount is related to the atrazine concentration. In the Step V, an amount of the secondary antibody (Ab2 ) is bounded
on the specific antibodies.
surface, and hence the impedance measured will change following
the changes of the concentration of the immobilized antigen, the
amount of the captured specific antibody, the competitive equilibrium between analyte, specific antibody and the competitor
antigen, as well as the amount of the captured secondary antibody
labelled with gold nanoparticles.
From all our measurements such changes are evident. From
impedimetric measurements, changes in the impedance of the
devices, following changes of the antigen concentration used for
the immobilization can be observed. In a similar way, changes are
also observed in the impedance related to the antibody dilution in
the incubation step. Adding atrazine during the competition step
implies that a fraction of this reagent will not be available and
hence, the captured amount will be lower than in the reference
sample. This change in the antibody concentration is a measure of
the used atrazine concentration.
A significant difference in the impedance spectra is observed
after the stepwise formation of the multilayer. Nyquist plots of
impedance spectra of layer-by-layer are shown in Fig. 3. This figure
shows an example of the difference in the impedance spectra of
the immunosensor functionalization, compared to the impedance
spectrum of the initial ID␮E. The goodness of the fitting between
the simulated and experimental spectra has been demonstrated
for all the curves from the sum of squares parameter. The sum of
squares is proportional to the average percentage error between
the original data points and the calculated values. This value was
between 0.01 and 0.03 for all cases. Fitting values of the ac equivalent circuit parameters are shown in Table 1.
As expected, the value of Rc is reasonably constant in all the
range of concentrations, and the CID␮E value shows small variations
related to the affinity binding of the biomaterial, which change the
electrical properties of the gap of the ID␮E.
Important changes can be seen, however, in the ohmic resistance
of the diluted PBS solution (Rs ) and in the polarization resistance
(Rp ), related to the covalent immobilization of the atrazine antigen
2d-BSA produced as are shown in Table 1. These changes are evident in the plots shown in Fig. 3, because Rs and Rp governing the
semicircles of the Nyquist plot [31]. Finally, the secondary antibody
capture produced an important change of Rs and a considerably
lower variation of Rp . At the same time a decrease in the Cdl was
observed.
3.3. Immunosensor conductive characterization
The physical interpretation of the different impedances related
to the final response is essential to understand and to interpret the
obtained results. The contribution of these impedances is described
below.
E. Valera et al. / Sensors and Actuators B 134 (2008) 95–103
99
Table 1
Fitting values of the equivalent circuit elements corresponding to the impedance spectra of Fig. 3 by commercially available software Zplot/Zview (Scibner Associates Inc.)
Elements of the equivalent circuit
ID␮E
Step I N-acetylcysteamine
Step II GPTS
Step III AT
0.5 ␮g mL−1
Step IV Ab1
0.25 ␮g mL−1
Step V Ab2 –gold
(40 nm) 1/1000
Rc ()
CID␮E (nF)
Rs ()
Rp ()
W − Q (nF(rad s−1 )1−(W-n) )
W−n
Cdl −T (nF(rad s−1 )1 −( Cdl −p) )
Cdl −p
104.9
131.1
6575
9023
10.622
0.98752
85.335
0.71886
94.12
135.3
9787
5595
30.857
0.97126
121.65
0.69894
106.2
133.8
7479
5292
34.895
0.97958
115.64
0.70737
115.9
136
12056
5660
13.544
0.90559
61.045
0.75873
113.5
135.9
14474
5715
11.22
0.89551
70.49
0.74527
114.4
135.5
17167
6202
11.438
0.89566
69.778
0.73778
W = 1/Q(ωj)n ; Cdl = 1/T(ωj)p ; where j is the imaginary unit and ω is the angular frequency.
3.3.1. Sensing structure
As it could be expected, the fact that there are gold nanoparticles
attached to a secondary antibody produces a large increase of the
current for the same voltage applied, and hence a large change in
the impedance element associated to this secondary antibody.
In order to make clearly explicit this part of the impedance
we split the impedance in two terms: the first is called blank
and includes the substrate impedance (Zboard ); the ohmic resistance of the electrolyte (Rs ); the contribution to the impedance of
the N-acetylcysteamine (ZN ); the contribution of the GPTS (ZGPTS );
the impedance of the antigen (ZAT ); and the impedance of the
primary antibody (ZAb1 ). The second term is the impedance of
interest named ZAb2 + gold particles , as it can be seen in Fig. 4. The
simplified structure shown in Fig. 4 is not a DC equivalent circuit; nevertheless it is useful to understand the dependences of
the measured impedance. This circuit is not a complete electrical
model of the device, as it does not include the effect of the capacitance between the electrodes, the double-layer capacitance and the
Warburg impedance.
3.4. Response of the atrazine immunosensor
In this work the atrazine levels in a solution are quantified by
conductive measurements, using the competitive antibody capture
system and the interdigitated electrodes described.
Fig. 3. Nyquist plot of impedance spectra corresponding to: (a) ID␮E, (b) Step I:
protection of ID␮Es with N-acetylcysteamine, (c) Step II: immunosensor surface
functionalization with GPTS, (d) Step III: antigen immobilization (0.5 ␮g mL−1 ),
(e) Step IV: competition step (specific primary antibody capture in 0.25 ␮g mL−1
concentration), and (f) Step V: capture of secondary antibody labelled with gold
(1/1000), taken in diluted PBS solution without redox couple. Symbols represent the
experimental data. Solid curves represent the computer fitting data with the parameters calculated by commercially available software Zplot/Zview (Scibner Associates
Inc.) using the equivalent circuit inset in the figure.
3.4.1. Conductive response
The quantitative tool that seems adequate to provide sensitivity
graphs is the measurement of the conductance after the secondary
labelled with gold antibody capture (Step V). The atrazine concentration should finally be related to the amount of gold nanoparticles
present on the immunosensor. In order to qualitatively show
how the immunosensor is sensitive to the atrazine concentration, experiments have been performed adding different values
of atrazine concentration (between 0.32 and 2000 ␮g L−1 ) during
the competition step (Step IV) using different ID’s arrays for every
concentration. Therefore, the amount of specific first antibody on
the ID␮E, and as consequence also the amount of Ab2 , is different
for each array. For these experiments 0.5 ␮g mL−1 concentration
of antigen immobilization (Step III), 0.25 ␮g mL−1 concentration of
specific primary antibody (Step IV), and 1/1000 concentration of
secondary labelled with gold antibody (Step V) were used. In addition, the blank (immunosensor impedance before Step V) was also
measured.
For the conductive measurements the electrodes were covered
by a diluted PBS solution with a conductivity of 1.6 ␮S cm−1 , and
these measurements were done using +25 and +100 mVdc bias.
DC voltages were chosen under 100 mV bias in order to avoid the
electrolysis of water. These values were extracted from sweep bias
voltage of +22.5 to +27.5 mVdc and +90 to +110 mVdc, respectively.
Conductive measurements and data processing are much simpler than the impedimetric measurements shown in Refs. [31–33].
Moreover, the fitting procedures to the equivalent circuit model, as
well as the parameter extraction process are completely avoided.
The results obtained can be seen in Fig. 5. In all cases the blank was
measured and reduced of the data measured after Step V in order to
Fig. 4. Sensing structure. The impedance of interest is the impedance related to the
secondary antibody, whereas the other impedance contributions are named blank.
100
E. Valera et al. / Sensors and Actuators B 134 (2008) 95–103
Fig. 5. Normalized calibration curves of the optimized atrazine immunoassay and the immunosensor presented: (a) immunosensor curves plotted do not include the blocking
step; (b) immunosensor curves plotted include the blocking step. Immunosensor curves for the atrazine detection related to the presence of gold particles (40 nm). Measures
were taken in diluted PBS solution and the blank was reduced. See Table 2 for the features of the optimized immunoassay and ID␮E immunosensor.
take into account only the contribution of the gold nanoparticles.
In the same way, the response to atrazine of the sensor follows an
inverse law and hence the response is larger at low concentrations
of atrazine. This is a result of the competitive method of detection
used in this work.
To validate the sensing approach, besides impedance spectroscopy measurements, our immunosensor has been characterized by means of chemical affinity methods. In Fig. 5 the results
obtained from the conductive measurements are compared with
the results of an ELISA assay on the same ID␮Es devices but using
a chemically raised colorimetric signal. In order to comparatively
show the immunosensor performance, in Fig. 5 the normalized
values of the change in the current as a function of the atrazine
concentration as well as the normalized results of the ELISA assay,
are plotted.
In Fig. 5a, the curves of immunosensor response for 25 and
100 mV bias voltage can be seen. The limits of detection (LODs)
obtained for the atrazine residues detection using the immunosensors shown in this work are in the order of 0.446 ␮g L−1 (100 mVdc
bias) and 1.217 ␮g L−1 (25 mVdc bias). Therefore, the LODs obtained
using the immunosensors developed are well below the MRL
required by EC for the atrazine in the wine grapes (100 ␮g L−1 ).
In Fig. 5b it can be seen the immunosensor response curves for
25 and 100 mV bias voltage but including a blocking agent (BSA
1% in PBS 30 min RT) after the covalent immobilization of the antigen (Step III). Blocking step was included to reduce non-specific
adsorption of Ab1 or Ab2 on the Pyrex.
The limits of detection obtained for the atrazine residues detection using the immunosensors presented in this work and including
the blocking agent are in the order of 0.104 ␮g L−1 (100 mVdc bias)
and 1.076 ␮g L−1 (25 mVdc bias). As in the case without blocking
agent, the LODs obtained are far below than the MRL required by
EC for the atrazine in the wine grapes.
A similar comparison is shown in Refs. [31,33]. As in these previous cases, ELISA assay is more sensitive than the immunosensor
described in this work, as a result of the ELISA high enzymatic
amplification of the chemical signal which is a phenomenon not
included in the interdigitated electrode sensor described here.
Nevertheless, the differences with the ELISA assay compared to
the results shown in Ref. [31] have been largely reduced in both
cases (using or not blocking agent), due to the optimizations
in the design and operation performed in this immunosensor,
such as the improvement of the adsorption technique, changing
the passive adsorption technique by the covalent immobilization,
and the inclusion of the secondary antibody labelled with gold
nanoparticles.
It is important to remark that the ELISA assay shows a similar
analytical profile that the one obtained from this immunosensor.
Therefore, we are confident that our immunosensor really reflects
the selective binding event. The more important analytical features
of the atrazine assays (conductimetric immunosensor and ELISA)
are shown in Table 2.
As it was explained above, impedimetric measurements were
taken in the 40 Hz to 1 MHz frequency range using a modulation
voltage of 25 mV amplitude. Therefore, in order to analyze the relation among the impedimetric and conductimetric measurements,
current value obtained at 40 Hz could be extrapolated to the current
value measured at 0 Hz using 25 mVdc bias (conductimetric measurement). For that, current values in a frequency range of 40 Hz
to 1 kHz were obtained from the simple division of the modulation voltage (25 mV) and the module of the impedance (|Z|). The
module of the impedance was chosen inset the real part of the
impedance in order to taking into account the impedance related
to the double-layer capacitance as well as the Warburg impedance.
In Fig. 6 are shown the current values of the (i) impedimetric
measurements between 40 Hz and 1 kHz; (ii) conductimetric measurements taken at 25 mVdc bias; (iii) trend which relates both
kinds of measurements.
3.5. Conductimetric immunosensor vs. impedimetric
immunosensor
As it has been commented above, previous works devoted to the
development of an impedimetric immunosensor for the atrazine
detection have been presented [31–33]. A comparison between
the conductimetric and impedimetric immunosensor performance,
when buffer is used as solution assay, is shown in Table 3.
Using the conductimetric immunosensor, very small LODs have
been obtained. These results are directly related to the secondary
antibody labelled with gold nanoparticles.
Despite the good results obtained with the conductimetric
immunosensor, we consider that its performance can still be
improved (in order to reach the LODs obtained fro the impedimetric
immunosensor) if two facts are taken into account:
First, it is important to remark that the label used for the conductimetric immunosensor is not an enzymatic label it is a conductive
label. Thus its influence is largely related to the aspect ratio between
the particle diameter (40 nm) and the electrodes gap (5000 nm). In
E. Valera et al. / Sensors and Actuators B 134 (2008) 95–103
101
Table 2
Features of the atrazine assaysa
Features of the atrazine assaysa
Conductimetric immunosensor
ELISA assay
Without blocking step
IC50 (␮g L−1 )
LOD (␮g L−1 )
R2
a
With blocking step
25 mV
100 mV
25 mV
100 mV
8.471 ± 0.19
1.217
0.89
5.293 ± 0.14
0.466
0.91
14.09 ± 0.25
1.076
0.82
11.54 ± 0.18
0.104
0.87
1.65
0.09
0.99
The parameters are extracted from the four-parameter equation used to fit the standard curve.
Table 3
Limit of detection of the atrazine assays performed in buffer solution (PBST)a
Features of the
atrazine assaysa
LOD (␮g L−1 )
a
Conductimetric immunosensor
(covalent immobilization)
0.104–1.217
Impedimetric
immunosensor
Passive adsorption
(wide frequency range)
Passive adsorption
(single frequency)
Covalent immobilization
(wide frequency range)
8.34
5.76–28.61
0.04
Limit of detection is extracted from the four-parameter equation used to fit the standard curve.
the conductimetric immunosensor presented here, this difference
is large and because of this the neighbourhood of each particle
becomes decisive, as it was studied previously [39]. Therefore, in
order to obtain a further improvement in the LOD this difference
must be reduced, for example reducing the electrodes gap.
Second, from the chemical point of view, one of the most
important differences that exist between both immunosensors is
that the conductimetric includes a second antibody. Inevitably,
the inclusion of the secondary antibody affects the immunosensor
performance. Thus, another interesting approach related the conductimetric immunosensor would be to directly include the gold
nanoparticle in the first antibody, therefore eliminating the second
antibody.
3.6. Red wine matrix effect
As it has been demonstrated above, the conductimetric
immunosensor shows trustworthy results in the case of the atrazine
detection. However the quantification of atrazine residues has been
performed in buffer solutions and not in real samples such a red
wine subject to matrix effects.
Here, first results of the red wine matrix effect found in our
immunosensors are described next. Red wine samples were chosen instead other matrixes such as white wine, water or grape
juice, because their strong matrix effect. Since if the red wine
Fig. 6. Impedimetric measurements extrapolated to DC conductimetric measurements. Impedimetric measurements were obtained from the simple division of the
modulation voltage (25 mV) and the module of the impedance (|Z|). Conductimetric
measurements were obtained at 25 mVdc bias.
matrix effect can be measured, the other matrix effects will be
easier.
Initially, non-treated red wine was used. Nevertheless, the
strong matrix effect of the red wine produced the inhibition of
the ELISA assay and also the inhibition of the changes in the standard impedimetric response. The undesired matrix effect in the
ELISA assay was avoided by means of a solid-phase extraction (SPE)
treatment, and diluting the extracts 1:50 with PBST. The same
protocol was followed for impedimetric measurements. Several
devices with standard functionalization and 0.5 ␮g mL−1 of antigen were treated with (i) red wine; (ii) specific antiserum at the
same concentration in buffer; (iii) red wine and specific antiserum
0.25 ␮g mL−1 .
From this experiment, some matrix effect exists which affects
both to the maximum signal of the assay as well as to the nonspecific signal. Nevertheless, these two effects do not produce the
total inhibition of the assay and therefore it is possible to evaluate
the concentration of atrazine in red wine.
4. Conclusions
A novel molecular selective conductimetric immunosensor for
the quantification of atrazine residues has been developed and fully
characterized. This immunosensor, which is mainly based on antibodies labelled with gold nanoparticles, has already demonstrated
high sensibility to atrazine, detecting atrazine concentrations
below the MRL required by EC for the atrazine in the wine
grapes.
Detection of the atrazine concentration is done by means of
simple DC measurements, thanks to the presence of the gold
nanoparticles which determines the conductance signal. Previously
to the inclusion of the labelled antibodies, a competitive reaction,
between the atrazine and the covalent immobilized antigen for a
small amount of the specific antibody, occurs on ID␮Es.
Although in this work, the characterization of the sensor has
been done using atrazine as the target pesticide, the concept of this
immunosensor can be easily applied for the detection of a broad
range of chemical or biological species if the appropriate antibody
and competitor are available.
The fabrication of the sensor described above is fast, simple
and inexpensive in mass production. Besides, thanks to the gold
nanoparticles, these sensors are more sensitive, the operation is
simple and does not require trained personnel or complex electronics.
102
E. Valera et al. / Sensors and Actuators B 134 (2008) 95–103
The use of a non-passivated interdigitated ␮-electrode increases
the device sensitivity to the changes on the chemical sensitive layer
at the cost of use a low conductivity buffer instead the conventional
isotonic buffers.
Although until now, the quantification of atrazine residues has
been done only in buffer, preliminary quantification of atrazine
residues in red wine are also described.
In the same way, the immunosensor characterization for other
compounds of interest in the food safety field and the development
of real analytical methodologies based on these devices are under
study in our laboratories and it will be the reported separately.
Acknowledgements
This work has been supported by the Ministry of Science and Technology (Contract numbers TEC2004-0121-E and
TEC2004-06854-C03-03/MIC) and by the European Community
(IST-2003-508774). The MNT group is a consolidated Grup de
Recerca de la Generalitat de Catalunya since the year 2001 (expedient 00329).The AMR group is a consolidated Grup de Recerca de
la Generalitat de Catalunya and has support from the Departament
d’Universitats, Recerca i Societat de la Informació la Generalitat de
Catalunya (expedient 2005SGR 00207).
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Biographies
Enrique Valera was born in Lima, Perú in 1977. He received
his bachelor’s degree in electronic engineering from the
Pontificia Universidad Católica del Perú in 2003. He joined
the Micro and Nanotechnologies group (MNTg) of Universitat Politècnica de Catalunya (Spain) in 2002. Mr.
Valera started to work toward his PhD in 2003. Since
2002 he is working in silicon micromachining devices,
gold nanoparticles dynamics, porous silicon technologies
for sensors development and biosensor technologies. His
main research areas are now biosensors, micro and nanotechnologies, nanoparticles manipulation, and MEMS.
Mr. Javier Ramón ([email protected]). PhD student working in Applied Molecular
Receptors Group. Chemistry degree (specialized in Organic Chemistry) by the University of Barcelona (Spain) in 2002. ERASMUS stay, experimental work in organic
chemistry by the University of Kaiserslautern (Germany) within 2001-2002. Title of
the work “Combinatorial palladium Catalyzed Cross Coupling Reactions” Director:
prof. Bernhard Witulski. Experimental master degree in Biological, Environmental
E. Valera et al. / Sensors and Actuators B 134 (2008) 95–103
and related Technologies by Instituto de Investigaciones Quı́micas y Ambientales (IIQAB) of CSIC within 2003–2004 with the title “Desarrollo de Técnicas
Inmunoquı́micas para Surfactantes Aniónicos de tipo LAS”. Director: Dra. Maria
Pilar Marco Colás. Since 2004 he has been working in AMR group leaded by Dr.
M.-Pilar Marco in IIQAB-CSIC (Spain) developing immunochemical protocols for
environmental monitoring and biosensors based in new transducing principles. This
involves experience in organic synthesis, chromatographic analysis, spectroscopic
characterization and molecular modeling of organic molecules, animal handling,
immunochemical techniques, surface chemistry, electrochemistry and development
of analytical methods. He is co-author of 5 publications of international relevance. The immunochemical techniques and immunosensors developed have found
application in different fields such as the clinical, food safety, environmental and
biological monitoring fields. He has participated in two European projects and one
Spanish project related with the biosensors and immunochemistry and he have an
I3P grant from Ministerio de Educación y Ciencia.
Dr. Francisco Sánchez-Baeza ([email protected], Staff Scientist, Group leader of
the AMRg group. PhD in chemistry by IQS (1987) and University of Barcelona (1991)
belongs to the IIQAB staff since year 1989 as Head of the Spectroscopy Services
and from 1998 as research scientist. He has been working in organic synthesis of
bioactive molecules and development of new organic synthetic reagents. He has
a long experience in spectroscopic techniques for identification and characterisation of organic compounds. This includes knowledge on the bases of spectroscopic
techniques and the application of computational chemistry to the determination of
structural and physico-chemical properties of small molecules. Since some years he
is staff scientist also co-directing projects and PhD students of the AMR group. His
actual research is focused in the development of selective molecular receptors, particularly the so-called artificial antibodies or molecular imprinted polymers (MIPs).
He has been co-author of more than 70 publications and the responsible scientist of
several EC and national projects.
Dr. M.-Pilar Marco ([email protected], Research Scientist, Head of the Applied
Molecular Receptors Group. PhD in pharmacy by the University of Barcelona in 1990.
103
Postdoctoral researcher at the University of California in Davis, developing immunochemical protocols for environmental and biological monitoring. Since 1993 she has
been working at the department of Biological Organic Chemistry of the IIQAB-CSIC.
She became part of the Scientific Research Staff of the CSIC on year 1996. She is leading and co-directing the AMR group composed by 8 PhD students, 2 post-docs and
a technical assistant. Her research is focused on the production of selective bioreceptors and the development of bioanalytical techniques. Particularly, she has been
very active on the production of specific antibodies for non-antigenic molecules. The
immunochemical techniques developed have found application in different fields
such as the clinical, food safety, environmental and biological monitoring fields. This
involves experience in organic synthesis, chromatographic analysis, spectroscopic
characterization and molecular modeling of organic molecules, animal handling,
immunochemical techniques and development of analytical methods. She has been
leading several EC and Spanish grants, directed several PhD theses and is a co-author
on more than 90 publications of international relevance. Her actual research interest are biosensors based in new transducing principles, evaluation and validation
of exposure and disease markers and investigation around mechanisms involved in
adverse drug reactions.
Ángel Rodrı́guez (M’96) graduated in telecommunication
engineering from the Universidad Politécnica de Cataluña
(Spain). From 1987 to 1992 he worked at IMEC (Belgium)
towards his doctoral work in the field of Polysilicon Thin
Film Transistors. In 1993, he became an Associate Professor in the Escuela Técnica Superior de Ingenieros de
Telecomunicación de Barcelona. Dr. Rodrı́guez has worked
in solar cells, Bipolar transistors, Polysilicon Thin Film
Transistors and MEMS. Concerning MEMS he has worked
in flow sensors, accelerators, different kinds of artificial
noses based in metallicoxides or polymers resonating
structures, RF MEMS, MEMS actuators, Porous Silicon
based MEMS/MOEMS and, more recently in Bio-MEMS.