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Trends in Analytical Chemistry, Vol. 27, No. 7, 2008
Impedance methods for
electrochemical sensors using
nanomaterials
Ian I. Suni
This article presents an overview of electrochemical sensors that employ
nanomaterials and utilize electrochemical impedance spectroscopy for
analyte detection. The most widely utilized nanomaterials in impedance
sensors are gold (Au) nanoparticles and carbon nanotubes (CNTs). Au nanoparticles have been employed in impedance sensors to form electrodes from
nanoparticle ensembles and to amplify impedance signals by forming
nanoparticle-biomolecule conjugates in the solution phase. CNTs have been
employed for impedance sensors within composite electrodes and as nanoelectrode arrays. The advantages of nanomaterials in impedance sensors
include increased sensor surface area, electrical conductivity and connectivity, chemical accessibility and electrocatalysis.
ª 2008 Elsevier Ltd. All rights reserved.
Keywords: Analyte detection; Biosensor; Carbon nanotube; Electrochemical impedance
spectroscopy; Electrochemical sensor; Gold nanoparticle; Immunosensor; Impedance;
Nanoparticle; Nanomaterial
Ian I. Suni*
Department of Chemical and
Biomolecular Engineering,
Center for Advanced Materials
Processing (CAMP), Clarkson
University, Potsdam, NY
13699-5705, USA
*
Tel.: +1 315 269 4471;
E-mail: isuni@clarkson.edu
1. Introduction
1.1. Electrochemical impedance
spectroscopy – background
Electrochemical impedance spectroscopy
(EIS) has long been employed for studying
electrochemical systems [1], including
those involved in corrosion, electrodeposition [2], batteries [3] and fuel cells [4].
For impedance measurements, a small
sinusoidal AC voltage probe (typically
2–10 mV) is applied, and the current
response is determined. The in-phase
current response determines the real
(resistive) component of the impedance,
while the out-of-phase current response
determines the imaginary (capacitive)
component. The AC probe voltage should
be small enough so that the system response is linear, allowing simple equivalent circuit analysis. Impedance methods
are quite powerful, in that they are capable of characterizing physicochemical
processes of widely differing time constants, sampling electron transfer at high
0165-9936/$ - see front matter ª 2008 Elsevier Ltd. All rights reserved. doi:10.1016/j.trac.2008.03.012
frequency and mass transfer at low frequency.
Impedance results are commonly fitted
to equivalent circuits of resistors and
capacitors, such as the Randles circuit
shown in Fig. 1 [5], which is often used to
interpret simple electrochemical systems.
This equivalent circuit yields the Nyquist
plot shown in Fig. 2, which provides visual
insight into the system dynamics. In Figs.
1 and 2, Rct is the charge-transfer resistance, which is inversely proportional to
the rate of electron transfer; Cd is the
double-layer capacitance; Rs is the solution-phase resistance; and, Zw is the
Warburg impedance, which arises from
mass-transfer limitations. If an analyte
affects one or more of these equivalent
circuit parameters and these parameters
are not affected by interfering species, then
impedance methods can be used for analyte detection. Rs arises primarily from the
electrolyte resistance and is analytically
useful mainly in conductivity sensors,
which I will not discuss in this article. The
Warburg impedance, which can be used to
measure effective diffusion coefficients, is
seldom useful for analytical applications.
The equivalent circuit elements in Figs.
1 and 2 that are most often useful for
analyte detection are Rct and Cd. The
measured capacitance usually arises from
the series combination of several elements,
such as analyte binding (Canal) to a sensing layer (Csens) on an Au electrode (CAu).
In this case, the measured capacitance is:
1
1
1
1
¼
þ
þ
Cd CAu Csens Canal
ð1Þ
for a sensing layer and analyte layer that
are continuous. In many cases, the
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Rct ¼ RAu þ Rsens þ Ranal
Cd
Rs
Zw
Rct
Figure 1. Randles equivalent circuit for a simple electrochemical
system.
-Z imag
Slope=unity
g
sin
w
a
cre
de
Rs
R ct
Z real
Figure 2. Nyquist plot arising from the Randles circuit shown in
Fig. 1.
capacitance at the Au electrode-sensing layer interface is
large and can be neglected. The sensitivity is then
determined by the relative capacitance of the analyte
layer and the sensing layer. For each dielectric layer, the
capacitance per unit area depends on the layer thickness
(t) according to:
C ed
¼
A
t
ð2Þ
where ed is the dielectric constant of the dielectric layer,
so capacitance is most sensitive to binding of large
analytes, such as proteins or cells.
One difficulty with capacitive sensors is that their
sensitivity depends on obtaining the proper thickness of
the original sensing layer [6]. If the original sensing
layer is too thin, then the underlying electrode surface
may be partially exposed, allowing for non-specific
interactions from interfering species. However, if the
original sensing layer is too thick, then the AC impedance current that is detected is dramatically reduced, as
is the change in capacitance upon analyte binding. Rct
can also be quite sensitive to analyte binding, particularly for detection of large species, such as proteins or
cells, which significantly impede electron transfer. For
analyte binding (Ranal) to a sensing layer (Rsens) on an
Au electrode (RAu), the measured resistance is:
ð3Þ
The resistance at the interface between the Au electrode
and the sensing layer is typically negligible. Measurement of Rct requires the presence of redox-active species
in the electrolyte. Impedance sensing is most useful for
large species that significantly perturb the sensing
interface, although impedance detection of glucose was
recently reported [7]. Many of the examples of impedance sensors that I discuss later in this article monitor
Rct as a measure of analyte concentration.
1.2. Electrochemical impedance spectroscopy – sensing
applications
For biosensors, EIS has some important advantages over
amperometry.
For direct amperometric biosensors, an oxidoreductase
enzyme is immobilized at a conductive electrode, and
electron transfer is detected during a biologically-mediated oxidation/reduction reaction. However, the active
site must be both in close proximity to the electrode
surface and easily accessible to the analyte solution. In
many cases, electron transfer occurs far from the electrode surface, and electron-transfer rates drop exponentially with distance [8]. This problem can be reduced
through the use of redox mediators, but detection then
becomes limited by mediator mass transfer.
Indirect amperometric biosensors detect the product of
a biologically-catalyzed reaction, often hydrogen peroxide. However, the analyte often contains additional
species (e.g., ureate or ascorbate) that can also be electrochemically oxidized or reduced, so indirect amperometric biosensors are not selective. One of the most
significant advantages of impedance detection for biosensing is that antibody-antigen binding can be directly
detected, allowing the development of immunosensors.
The main drawback of impedance methods for biosensors is the need for interfacial engineering to reduce
non-specific adsorption. One well-studied method to
minimize non-specific interactions is to embed the probe
agent into a composite film that contains the biomolecule of interest interspersed with a protein-resistant
species, such as molecules containing ethylene-glycol
moieties. This approach has been widely touted by the
research group of George Whitesides [9–11], and such
reagents are now commercially available. The use of
impedance methods for biosensors has been recently
reviewed [12], but not with a focus on the use of
nanomaterials. Limits of detection (LODs) have been reported for impedance biosensors in the nM–pM range in
controlled laboratory conditions [13–17].
It should be acknowledged that Au-nanoparticle
conjugation to biomolecules has been employed in biosensors using several other electrochemical detection
methods, predominantly anodic stripping voltammetry
(ASV), anodic Au surface oxidation and quartz crystal
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microbalance (QCM) [18,19]. However, impedance
detection has some significant advantages over these
methods. Impedance sensing does not require the voltage scanning needed for ASV and anodic oxidation,
which is time consuming and may degrade the electrochemical interface during wide potential sweeps. In
addition, impedance methods are largely insensitive to
environmental disturbance, which is often problematic
for QCM sensors.
2. Nanomaterials for sensing applications
Nanomaterials are generally defined as involving the
length scale from 1–100 nm; in other words, materials
intermediate between the atomic and molecular scale
and the bulk scale. The chemical, electronic, and optical
properties of nanomaterials generally depend on both
their dimensions and their morphology. Although a wide
variety of nanomaterials for sensors have been reported
in the literature, the most widely employed nanomaterials are carbon nanotubes (CNTs) and Au nanoparticles,
in part because of their commercial availability. In
addition, both materials are considered to be biocompatible.
2.1. Au nanoparticles
Au nanoparticles are generally synthesized by chemical
reduction of Au salts in aqueous-phase, organic-phase,
or mixed-phase solutions [20]. The most difficult aspect
of this synthesis is to control the reactivity of the Aunanoparticle surface during particle growth, since the
surface energy is quite high. Controlled synthesis of Au
nanoparticles requires the use of stabilizing agents, such
as citrate or thiolated species, that bind to the particle
surface to control growth and to prevent aggregation.
Numerous methods have been reported for creation of
biomolecule-Au-nanoparticle conjugates either during
or after Au-nanoparticle synthesis [20]. Commercial reagents are now available for conjugation of biomolecules
to Au nanoparticles of several different sizes. One of the
primary reasons for the intensive research into biomolecule-Au-nanoparticle conjugates is that biomolecules in
this environment are generally stable and retain their
biological activity. Depending on the application, different Au-nanoparticle sizes may be optimal [21].
2.2. Carbon nanotubes
CNTs, which are allotropes of carbon from the fullerene
structural family, can be conceived as sp2 carbon atoms
arranged in grapheme sheets that have been rolled up into
hollow tubes. Multi-walled CNTs (MWCNTs) behave as
conductors and have electrical conductivities greater than
metals, suggesting their incorporation into sensing electrodes may be beneficial. However, depending on the tube
diameter and chirality, single-walled CNTs (SWCNTs) can
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behave electronically as either metals or semiconductors
[22], complicating their use in sensing electrodes. CNTsynthesis methods create a mixture that includes amorphous carbon, graphite particles and CNTs, so synthesis is
typically followed by a difficult separation process.
For electrochemical applications, CNTs are typically
activated in strong acids, which opens the CNT ends and
forms oxygenated species, making the ends hydrophilic
and increasing the aqueous solubility of CNTs [22]. The
electrochemical behavior of CNTs varies considerably
with the methods used for purification and preparation,
including oxidation treatment [22]. For analytical
applications, CNTs are most often used to modify other
electrode materials, or as part of a composite electrode,
in part due to difficulties in handling them.
3. Impedance sensors using Au nanoparticles
3.1. Au-nanoparticle substrates – impedance detection
The most widely reported use of Au nanoparticles in
impedance sensors involves their incorporation into an
ensemble substrate onto which a protein, oligonucleotide, or other probe molecule is immobilized [23–34].
Most studies involve construction of a sensing interface
that contains one layer of Au nanoparticles on a conductive electrode, although, in a few cases, Au nanoparticles are incorporated into a ceramic sol gel or
polymer film. The Au nanoparticles are sometimes made
using colloidal techniques, and sometimes by electrodeposition. The Au nanoparticles can be conjugated with
probe reagents (antibodies or ssDNA) either before or
after the Au-nanoparticle ensemble is formed.
The advantages of sensing interfaces that contain Aunanoparticle networks, compared to sensing interfaces
based on flat Au surfaces, include the increased surface
area for sensing, improved electrical connectivity
through the Au-nanoparticle network, and chemical
accessibility to the analyte through these networks. The
advantages, compared to non-Au surfaces, also include
electrocatalysis.
One potentially powerful method for using Au nanoparticles to enhance impedance detection in biosensors
involves the construction of three-dimensional networks
with Au nanoparticles dispersed throughout the sensing
interface. This can be accomplished through repeated
use of a bifunctional reagent, such as cysteineamine or
4-aminothiophenol, where the thiol group can bind to a
biomolecule and the amine group can bind to Au
nanoparticles, for layer-by-layer formation of an Aunanoparticle network. Impedance detection of human
immunoglobulin (hIgG) using such a three-dimensional
Au-nanoparticle network was recently reported using 6nm diameter Au nanoparticles and cysteamine as the
bifunctional reagent [23]. Fig. 3 shows the sensorpreparation process.
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Figure 3. Au-nanoparticle-multilayer preparation onto an Au electrode (a), and the immobilization of antibody and the interaction of antigen and
biotin-conjugated antibody (b) (from [23]).
These authors also studied the nature of the sensing
interface as a function of the number of Au-nanoparticle
layers using both cyclic voltammetry and EIS. As the
number of Au-nanoparticle layers increased, the
3=4
FeðCNÞ6
oxidation/reduction peak height increased
and the peak separation decreased, demonstrating increased reversibility. Similarly, Rct decreased continuously as the number of Au-nanoparticle layers increased.
Following electrostatic binding of goat anti-human
IgG antibody, the sensing interface was able to detect the
presence of hIgG, with Rct increasing with an increase in
human-IgG concentration. Amplification of the impedance signal was accomplished by further binding biotinconjugated goat anti-human IgG, resulting in a detection
range of 5–400 lg/L. The LOD was then estimated to be
0.5 lg/L [23]. In this study, antibody was immobilized
only on the outer layer of Au nanoparticles to ensure
chemical accessibility of the analyte, a protein. For
small-molecule analytes that can be detected by impedance methods, such multilayer Au-nanoparticle net-
works may be invaluable for sensing. They could allow
dramatic increases in the electrode surface area without
introducing mass-transfer limitation.
Impedance detection of carcinoembryonic antigen
(CEA), a glycoprotein involved in cell adhesion produced
only during fetal development, was recently reported
[31]. The CEA antibody was first bound through its
surface amino groups to glutathione-modified Au
nanoparticles of diameter 15 ± 1.5 nm by amide-bond
formation using N-(3-dimethylaminopropyl)-N 0 -ethylcarbodiimide hydrochloride (EDC) and N-hydroxysulfylsuccinimide sodium salt (NHSS). The sensing interface
was then formed by co-polymerizing a mixture of
o-aminophenol and the Au-nanoparticle-conjugated
CEA antibodies. An interesting feature of their study was
the direct comparison between the antibody-containing
sensing interface, with and without Au-nanoparticle
conjugation. They reported that Rct increased by only
0.59 · 105 X (35%) for the sensing interface without
Au nanoparticles, but by 6.3 · 105 X (7%) with Au
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nanoparticles [31]. The authors tested their sensing
interface in both model lysozyme solutions and serum
samples and reported no false positives arising from
non-specific interactions. They estimated an LOD for CEA
of 0.1 ng/mL.
Impedance detection was recently demonstrated for an
intriguing application, detection of the IgE antibody to a
protein allergen from dust mites [29,30]. Au nanoparticles were deposited onto a glassy-carbon electrode
(GCE) either by electrodeposition, or by immersion in
(3-mercaptopropyl)trimethoxysilane (MPTS), followed
by immersion in a colloid solution containing 16-nm
diameter Au nanoparticles. For Au electrodeposition,
30 s of deposition from 0.1% HAuCl4 produced Au
nanoparticles of average diameter 40 ± 8 nm. The Aunanoparticle-modified GCE was then immersed in recombinant dust-mite allergen Der f2 to form a protein
film, and this interface was employed for impedance
detection of the murine monoclonal antibody to Der F2
over the range 2–300 lg/ml. At relatively low antibody
concentrations, Rct increased continuously with antibody concentration. At higher antibody concentrations,
Rct became relatively insensitive to changes in antibody
concentration, probably due to surface saturation. This
type of impedance sensor might be employed for allergy
screening of patients, where allergen-specific IgE is detected for a wide range of allergens.
In addition to using antibodies as probe reagents,
impedance sensors have been demonstrated using DNA
or oligonucleotides bound to Au-nanoparticle arrays to
detect complementary target molecules [26,34]. In
addition, the incorporation of CdS nanoparticles conjugated to ssDNA into the sensing interface of an impedance sensor has been reported [35]. One group reports
forming a sensing interface by binding thiol-derivatized
oligonucleotides onto Au surfaces modified by Au electrodeposition, followed by impedance detection of two
minor DNA groove-binding agents, mythramycin and
netropsin, and a DNA intercalator, nogalamycin [26].
The advantage of using Au electrodeposition to modify
the Au substrate is that the effects of surface roughness,
which is related to the Au-nanoparticle size, can be
studied quantitatively by measuring the surface area by
voltammetric reduction of Au oxide. Substrates were
prepared with a total surface area up to 90% greater
than that of the original flat Au substrate. The greatest
sensitivity was observed for an Au-electrodeposition
process that produced Au nanoparticles in the 20–80nm range. The authors estimated that this allowed a
reduction in the LOD by a factor of 20–40x, down to
5 nM for nogalamycin [26].
Au nanoparticles and carbon nanofibers have also
been reported to be useful in composite substrates for
impedance sensing of cells [36,37]. In these studies
using EIS, the binding of K562 leukemia cells was
monitored as an increase in Rct. These authors reported
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that incorporating Au nanoparticles increased the sensitivity to cell binding, which was attributed to increased
electrode-surface area. Au nanoparticles were first synthesized using chitosan as a combined reducing and
stabilizing agent, then reacted with ammonia to create a
sol-gel film atop a GCE with embedded Au nanoparticles
of 12-nm diameter. Adhesion of K562 leukemia cells
was then monitored in situ by EIS. Cell adhesion could be
detected only by the combination of chitosan and Au
nanoparticles atop a GCE. Rct was reported to correlate
to the logarithm of the cell concentration over the range
104–108 cells/mL with an LOD of 8.7 · 102 cells/mL.
3.2. Au-nanoparticle conjugation in solution
Several recent studies described different strategies for
the use of Au nanoparticles for impedance sensing that
involved Au-nanoparticle conjugation in the solution
phase rather than prior modification of the sensing
interface.
In one approach, impedance sensing included an extra
step of analyte conjugation to 10-nm diameter Au
nanoparticles, with signal amplification occurring only
when the Au nanoparticles become embedded in the
sensing interface [38]. This approach was demonstrated
using the model system fluorescein/anti-fluorescein,
with fluorescein bound to the flat Au substrate using
EDC/NHSS linker chemistry. The analyte (goat antifluorescein) was conjugated to Au nanoparticles in
solution prior to detection. A change in the impedance at
the sensing interface was observed only when the antibody was conjugated to Au nanoparticles, but not for the
bare antibody [38]. Signal amplification was significantly higher with a redox probe (impedance detection)
than without a redox probe (capacitance detection). This
is believed to reflect the substantial electrochemistry that
can occur on the Au nanoparticles embedded within the
sensing interface, which is otherwise essentially a polymer film. As a result, Rct is substantially reduced upon
analyte binding, which embeds Au nanoparticles within
the sensing interface.
A similar detection scheme was recently reported to
detect DNA hybridization, with the target ssDNA conjugated to 5-nm diameter CdS nanoparticles [39]. Probe
ssDNA was immobilized onto an Au electrode using selfassembly chemistry and amide-bond formation with
EDC/NHSS coupling. CdS nanoparticles were prepared
by precipitation from CdCl2 and Na2S using mercaptoacetic acid as a stabilizer, then conjugated to the
complementary ssDNA. The authors reported that conjugation to CdS nanoparticles increased the sensitivity by
about two orders of magnitude. Interestingly, unlike the
results observed with Au-nanoparticle conjugation [38],
here analyte binding was accompanied by a dramatic
increase in Rct [39]. The difference between these two
studies can be explained by the different rates of electron
transfer on Au and CdS, and by the different sensing
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interfaces. For CdS-nanoparticle conjugation, the interfacial Rct prior to analyte detection was about two orders
of magnitude lower than that in the study with Aunanoparticle conjugation. For this less well-passivated
sensing interface, the dominant effect upon binding of
ssDNA-CdS is obscuration of the underlying conductive
electrode, rather than enhanced rates of electron transfer
due to embedding of CdS nanoparticles. However, when
Au nanoparticles are embedded into a sensing interface
that is completely polymer coated, the dominant effect is
the improved rates of electron transfer on the Au
nanoparticles.
In biosensors, the use of nanomaterials has been
envisioned to create successive amplification steps [40].
This type of approach was recently demonstrated with a
different type of solution-phase Au-nanoparticle conjugation, utilizing a strategy that might be termed an
impedance-sandwich assay [41]. In this approach, antiprotein A IgG was bound to an Au-electrode surface, and
then exposed to protein A of varying concentrations.
Following protein A binding, the sensing interface was
exposed to a solution containing IgG antibodies conjugated to 13-nm diameter Au nanoparticles. Without this
amplification step, the LOD of protein A was reported to
be 1.0 ng/mL, and the LOD was reduced by one order of
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magnitude by the amplification step. The authors reported that their sensitivity was about 100x better than
that obtained with conventional ELISAs. One advantage
of this approach is that the protein-antibody conjugate
can be prepared in advance and stored for about one
month without loss of activity.
Another group recently reported the use of solutionphase Au-nanoparticle conjugation for amplifying the
signal from an impedance biosensor. The sensing interface was an Au electrode onto which Au nanoparticles
were attached using 1,6-hexanedithiol, followed by
immobilization of rabbit anti-IgG [28]. After binding the
hIgG analyte, and blocking non-reacted surface sites
with bovine serum albumin (BSA), the impedance signal
was amplified by binding Au-colloid-labeled goat antihIgG that was synthesized in advance [28]. This approach (Fig. 4) was motivated by the relatively small
impedance change sometimes observed upon antigen
recognition by a surface-immobilized antibody. Without
amplification, the impedance change upon binding of
hIgG was about 100 X-cm2, whereas, with amplification, the impedance change was several thousand
X-cm2. The authors reported an LOD for human IgG of
4.1 ng/L and a linear concentration range of about
15–330 ng/L.
Figure 4. The process of immobilization of rabbit anti-hIgG antibody onto an Au electrode, followed by analyte binding and amplification by the
Au-nanoparticle-labeled antibody (from [28]).
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4. Impedance sensors using carbon nanotubes
4.1. Carbon-nanotube substrates – impedance detection
The most detailed studies of impedance sensors that
employ CNTs do not employ SWCNTs or MWCNTs, but
instead employ electrodes constructed from CNT towers
grown by chemical-vapor deposition (CVD) [42–44].
Starting with bare Si wafers, Al was deposited by electron-beam evaporation and then oxidized, followed by
deposition of a Fe-catalyst film through a shadow mask.
CNT towers several mm thick were then grown by CVD
at 750C from a mixture of ethylene, water, and
hydrogen. The CNT tower was peeled from the Si substrate, cast in epoxy, and polished to reveal the underlying CNTs. The average CNT diameter is 20 nm, the
average spacing is about 200 nm, and the aspect ratio is
approximately 2 · 105. A significant advantage of this
method for creating an electrochemical-sensing interface
is that purification of the CNTs is not needed.
The electrochemical characteristics of these CNTtower electrodes have been most fully characterized by
voltammetry. Voltammetry of CNT towers in both
FeðCNÞ63=4 and RuðNH3 Þ3þ
show a sigmoidal shape,
6
without clear current peaks, at scan rates of up to
100 mV/s, and show current peaks for scan rates of
500 mV/s and above [42,43]. These results are similar
to results for MWCNT arrays that exhibit sigmoidal
voltammograms for large nanotube spacing, where the
diffusion fields from individual nanotubes do not fully
overlap, and peak-shaped voltammograms for small
nanotube spacing, where diffusion fields overlap [45,46].
As has been widely reported for micro-electrodes [47],
arrays of nanotube electrodes have enhanced diffusion
rates relative to macroscopic electrodes, and reduced
capacitance per unit area, which can significantly improve their sensitivity. Given the high electron-transfer
rates observed, CNT towers might be useful for characterizing rapid redox processes [42].
CNT-tower electrodes have been employed for
impedance detection of both mouse IgG and prostatecancer cells [43,44]. Prior to immobilization of antimouse IgG, the open end of the CNTs were oxidized in
strong acid or electrochemically to form active carboxylate groups [43]. This allowed the use of standard EDC/
NHSS coupling chemistry for amide-bond formation to
anti-mouse IgG. Both antibody immobilization and analyte binding were monitored by the extent to which they
increased Rct, providing a non-linear calibration curve.
The LOD for mouse IgG was reported as 200 ng/mL, with
a dynamic range of up to 100 lg/mL. Preliminary results
for impedance detection of prostrate-cancer cells involved
somewhat more complex electrode preparation, including
Au electrodeposition onto the CNT-tower electrode [44].
As for protein detection, cell binding is detected as an
increase in Rct.
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CNTs have also been incorporated into composite
electrodes used for impedance detection of DNA hybridization [48,49]. In these studies, MWCNTs were copolymerized with polypyrrole atop a GCE. EDC/NHSS
linker chemistry was used to form an amide bond and
immobilize ssDNA. The complementary oligonucleotide
could be detected by the accompanying change in Rct,
both with [48] and without [49] subsequent metallization. DNA metallization is a widely-studied technique,
whereby metal ions that bind to the center of the DNA
double helix greatly increase the conductivity of the
sensing interface, and that could be detected as a
reduction in Rct [48].
However, DNA hybridization without metallization
could be detected as an increase in Rct [49]. CNTs were
incorporated within the sensing interface due to their
high conductivity and their effect of increasing the active
surface area.
5. Future outlook
EIS has been widely used to study a variety of other
electrochemical systems, including corrosion, electrodeposition, batteries and fuel cells. However, only
recently have impedance methods been applied in the
field of biosensors. Given their ability to sense Rct and Cd,
application should be possible for several different types
of sensing schemes, with numerous recognition agents.
Electrochemical impedance sensors are particularly
promising for portable and implantable applications.
Commercialization will depend on improvements in
several different areas, including minimization of the
effects of non-specific adsorption.
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