Impedance methods for electrochemical sensors using nanomaterials

Trends 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 

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 

¼ 

Cd 

1 

þ 

CAu 

1 

þ 

Csens 

1 

ð1Þ 

Canal 

for a sensing layer and analyte layer that 

are continuous. In many cases, the 

0165-9936/$ - see front matter ª 2008 Elsevier Ltd. All rights reserved. doi:10.1016/j.trac.2008.03.012 604

Trends in Analytical Chemistry, Vol. 27, No. 7, 2008 Trends 

-Z imag 

R s 

R s 

decreasing w 

R ct 

R ct 

Z real 

C d 

Z w 

Figure 1. Randles equivalent circuit for a simple electrochemical 

system. 

Slope=unity 

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 ð2Þ 

A t 

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) onan 

Au electrode (RAu), the measured resistance is: 

Rct ¼ RAu þ Rsens þ Ranal 

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 

http://www.elsevier.com/locate/trac 605


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 sp 2 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 

606 http://www.elsevier.com/locate/trac 

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.


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 

FeðCNÞ 

3 =4 

6 

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 · 10 5 X (35%) for the sensing interface without 

Au nanoparticles, but by 6.3 · 10 5 X (7%) with Au 



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 


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. R ct was reported to correlate 

to the logarithm of the cell concentration over the range 

10 4 –10 8 cells/mL with an LOD of 8.7 · 10 2 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


interfaces. For CdS-nanoparticle conjugation, the interfacial 

R ct 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 

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-cm 2 , whereas, with amplification, 

the impedance change was several thousand 

X-cm 2 . 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]). 



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 750°C 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 · 10 5 . 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 

3 =4 

FeðCNÞ6 and RuðNH3Þ 3þ 

6 show a sigmoidal shape, 

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. 


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 R ct and C d, 

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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