10.1039@b714449k

Electrochemical biosensors

Niina J. Ronkainen,*aH. Brian Halsall

band William R. Heineman

b

Received 3rd November 2008

First published as an Advance Article on the web 1st February 2010

DOI: 10.1039/b714449k

Electrochemical biosensors combine the sensitivity of electroanalytical methods with the inherent

bioselectivity of the biological component. The biological component in the sensor recognizes its

analyte resulting in a catalytic or binding event that ultimately produces an electrical signal

monitored by a transducer that is proportional to analyte concentration. Some of these sensor

devices have reached the commercial stage and are routinely used in clinical, environmental,

industrial, and agricultural applications. The two classes of electrochemical biosensors,

biocatalytic devices and affinity sensors, will be discussed in this critical review to provide an

accessible introduction to electrochemical biosensors for any scientist (110 references).

1. Introduction

1.1 Background

Sensors are devices that register a physical, chemical, or bio-

logical change and convert that into a measurable signal.1 The

sensor contains a recognition element that enables the selective

response to a particular analyte or a group of analytes, thus

minimizing interferences from other sample components

(Fig. 1). Another main component of a sensor is the transducer

or the detector device that produces a signal. A signal processor

collects, amplifies, and displays the signal.

Electrochemical biosensors, a subclass of chemical sensors,

combine the sensitivity, as indicated by low detection limits, of

electrochemical transducers with the high specificity of bio-

logical recognition processes. These devices contain a biological

recognition element (enzymes, proteins, antibodies, nucleic

acids, cells, tissues or receptors) that selectively reacts with

the target analyte and produces an electrical signal that is

related to the concentration of the analyte being studied.

Electrochemical biosensors can be divided into two main

categories based on the nature of the biological recognition

process i.e. biocatalytic devices and affinity sensors.2 Bio-

catalytic devices incorporate enzymes, whole cells or tissue

slices that recognize the target analyte and produce electro-

active species. Special emphasis will be placed on enzyme

electrodes for the detection of glucose, lactose, and xanthine.

Affinity sensors rely on a selective binding interaction between

the analyte and a biological component such as an antibody,

Fig. 1 A schematic of a biosensor with electrochemical transducer.

aDepartment of Chemistry, Benedictine University, 5700 College Road,Lisle, IL 60532-0900, USA. E-mail: [email protected];Fax: +1 630 829 6547; Tel: +1 630 829 6549

bDepartment of Chemistry, University of Cincinnati,P.O. Box 210172, Cincinnati, OH 45221-0172, USA.E-mail: [email protected], [email protected];Fax: +1 513 556 9239; Tel: +1 513 556 9274, +1 513 556 9210

Niina J. Ronkainen

Niina J. Ronkainen receivedher BS in chemistry andbiology at Butler University(Indianapolis, USA) in 1997and her PhD at the Universityof Cincinnati (USA) in 2003where she specialized in bio-analytical chemistry. From2003–2004 she taught chemis-try as a visiting assistantprofessor at Tulane University(New Orleans, USA). In 2004she joined Benedictine Univer-sity as an assistant professorof chemistry. She currentlydoes basic research in bio-

sensors and electrochemistry. She is an active member of theChemical Education division of the American Chemical Society.

H. Brian Halsall

H. Brian Halsall is a professorof chemistry, and a member ofthe Sensors & Biosensors Groupin the Department of Chemistryat the University of Cincinnati.He received a BSc (Hons) andPhD in chemistry at the Uni-versity of Birmingham, UK.This was followed by post-doctoral work at UCLA, afterwhich he joined the staff of theMAN Program at Oak RidgeNational Laboratory beforesettling in Cincinnati. Hisprincipal research interestsinclude biosensors, electro-chemical immunoassay, andglycoprotein biochemistry.

This journal is �c The Royal Society of Chemistry 2010 Chem. Soc. Rev., 2010, 39, 1747–1763 | 1747

CRITICAL REVIEW www.rsc.org/csr | Chemical Society Reviews

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nucleic acid, or a receptor. Immunosensors and DNA hybridi-

zation biosensors with electrochemical detection will be

discussed as examples of affinity sensors.

Biosensors constitute an interdisciplinary field that is

currently one of the most active areas of research in analytical

chemistry. Using biosensors typically eliminates the need for

sample preparation. The biosensor’s performance is usually

experimentally evaluated based on its sensitivity, limit of

detection (LOD), linear and dynamic ranges, reproducibility

or precision of the response, selectivity and its response to

interferences.1 Other parameters that are often compared

include the sensor’s response time (i.e. the time after adding

the analyte for the sensor response to reach 95% of its final

value), operational and storage stability, ease of use and

portability. Ideally, the sensing surface should be regenerable

in order for several consecutive measurements to be made. For

many clinical, food, environmental, and national defense

applications, the sensor should be capable of continuously

monitoring the analyte on-line. However, disposable, single-use

biosensors are satisfactory for some important applications

such as personal blood glucose monitoring by diabetics.

1.1.1 Biocatalytic sensors. Although many types of bio-

recognition elements have been used in biosensing devices,

electrochemical biosensors primarily use enzymes due to their

high biocatalytic activity and specificity.3 Biocatalytic sensors

using enzymes as the recognition element often have relatively

simple designs and do not require expensive instrumentation.

Such sensors are typically easy to use, compact, and inexpensive

devices. Different detection configurations can be used such as

stationary sample solution vs. flow conditions or bulk sample

solution vs. a microdrop detected using a microelectrode.

Biocatalytic sensors can also be easily adapted to automatic

clinical lab and/or industrial analysis. Personal blood glucose

monitoring devices are the most successful commercial applica-

tion of biocatalytic sensors.

Biocatalytic sensors incorporate biological components

such as enzymes, whole cells or tissue slices that recognize

the target analyte and produce electroactive species or some

other detectable outcome.2 Enzymes, globular proteins com-

posed mainly of the 20 naturally occurring amino acids that

catalyze biochemical reactions, are the oldest and still most

commonly used biorecognition element in biosensors.2,4

Enzymes can increase the rate of a reaction significantly

relative to an uncatalyzed reaction. The enzyme–substrate

interactions can be characterized by kinetic studies. Para-

meters such as origin and availability of the biological com-

ponent, its operational and storage stability as well as

immobilization procedure should be considered when preparing

a biocatalytic sensor.5 Also, sensitivity of the biorecognition

element to experimental conditions such as pH, temperature,

and stirring should be minimal and variation between measure-

ments should be as low as possible.3 Because of their complex

molecular structures, enzymes often have exquisite specificity

for their substrate molecule and can detect individual sub-

stances in a complex mixture, such as urine or blood, very

selectively. This removes the need for time-consuming, labor-

intensive, and interference-prone sample pretreatment and

separation steps used in composite methods. The arrangement

of amino acids at the active site of the enzyme, often found at

the centroid of the protein, bind with the specific sub-

strate making the enzyme selective for one type of substrate

molecule.4 Many enzymes also incorporate small nonprotein

chemical groups, such as cofactors or prosthetic groups, into

the structures of their active site that help determine substrate

specificity.4 The inherent selectivity of enzymes often circumvents

the signals produced by interfering species that are sometimes

found in complex samples. However, enzyme activity is often

further modulated by other components such as activators and

inhibitors.4 Researchers also had to find ways to manage the

enzyme adsorption that could lead to electrode fouling as well

as denaturation and loss of enzyme’s catalytic activity on the

electrode surface.5 Biocatalytic biosensors will be described in

more detail in Section 2.

Many biochemical analytes of interest are not amenable to

detection by enzyme electrodes due to the lack of sufficiently

selective enzymes being available for the analyte or the analyte

not being commonly found in living systems.1,5 That is when

affinity biosensors are considered as an alternative method.

1.1.2 Affinity biosensors. Affinity sensors use the selective

and strong binding of biomolecules such as antibodies (Ab),

membrane receptors, or oligonucleotides, with a target analyte

to produce a measurable electrical signal.2 The molecular

recognition in affinity biosensors is mainly determined by

the complementary size and shape of the binding site to the

analyte of interest.2 The high affinity and specificity of the

biomolecule for its ligand make these sensors very sensitive

and selective.1 The binding process such as DNA hybridization

or antibody–antigen (Ab–Ag) complexation is governed by

thermodynamic considerations.2

Immunosensors are Ab-based affinity biosensors where the

detection of an analyte, an antigen or hapten, is brought about

by its binding to a region of an Ab.6 The electrochemical

transducer responds to the binding event and converts the

electrical response to an output that can be amplified, stored,

and displayed. Complementary regions of the Ab bind to an

Ag that was used to produce the antibodies in a host organism

William R. Heineman

William R. Heineman is a Dis-tinguished Research Professorin the Department of Chemis-try at the University ofCincinnati. He received a BSin chemistry at Texas TechUniversity and a PhD at theUniversity of North Carolinain Chapel Hill and was a post-doctoral associate at CaseWestern Reserve Universityand The Ohio State Univer-sity. His research interestsinclude spectroelectrochemistry,electrochemical immunoassay,sensors, and bioanalytical

chemistry. He is a recipient of the Charles N. Reilley Awardin Electroanalytical Chemistry and the Torbern Bergman Medalfrom the Analytical Section of the Swedish Chemical Society.

1748 | Chem. Soc. Rev., 2010, 39, 1747–1763 This journal is �c The Royal Society of Chemistry 2010

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such as a rabbit or a mouse with high specificity and affinity.4

Such polyclonal Abs are heterogeneous with respect to their

binding domain, and may be refined by a selection process to

yield monoclonal Abs—MAbs—all of whose members of a

particular MAb clone are identical. Abs and MAbs can be

developed for a wide range of substances. Theoretically, if an

Ab can be raised against a particular analyte, an immuno-

sensor could be developed to detect for that substance. Immuno-

sensors are well known among analytical methods for their

extremely low detection limits.6 Immunoassays and immuno-

sensors have been developed for both quantitative and quali-

tative applications.1,2 Immunosensors can be used to detect

trace levels (ppb, ppt) of bacteria, viruses, drugs, hormones,

pesticides, and numerous other chemicals.1,2 Examples of

immunosensor applications include monitoring food safety

related to severe allergies (such as peanuts), detecting environ-

mental pollutants such as herbicides and pesticides in water

and soil, detecting biomedical substances such as warfarin,

and monitoring for biowarfare agents such as toxins, bacteria,

viruses, and spores.1,2 Relatively inexpensive kits such as for

home pregnancy and fertility tests can be produced once the

assay is fully developed. In the past, the limited availability of

Ab varieties mainly produced by university and small bio-

technology companies has slowed down the affinity biosensor

development.6 However, antibodies are now sold by many

sources including large manufacturers of laboratory reagents

such as Sigma Aldrich.

Nucleic acids have been less commonly used as the bio-

recognition element in affinity sensors compared to antibodies.

Biorecognition using DNA or RNA nucleic acid fragments

relies on either complementary base-pairing between the sensor’s

nucleic acid sequence and the analyte of interest, or generating

nucleic acid structures, known as aptamers, that recognize and

bind to three-dimensional surfaces, such as those of proteins.

Nucleic acids are now becoming of greater importance as the

biorecognition agent in sensors since a recent rapid expansion

in knowledge of their structure and how to manipulate them.1

DNA affinity probes are typically used in medical diagnostics

to detect cancers, viral infections, and genetic diseases.1 Affinity

biosensors will be described in more detail in Section 3.

1.2 Electrochemical detection

Most biosensors use electrochemical detection for the trans-

ducer because of the low cost, ease of use, portability, and

simplicity of construction.1,2 The reaction being monitored

electrochemically typically generates a measurable current

(amperometry), a measurable charge accumulation or poten-

tial (potentiometry) or alters the conductive properties of the

medium between electrodes (conductometry).3 Use of electro-

chemical impedance spectroscopy by monitoring both resis-

tance and reactance in the biosensor is also becoming more

common.3

Electrochemistry is a surface technique and offers certain

advantages for detection in biosensors. It does not depend

strongly on the reaction volume, and very small sample volumes

can be used for measurement.6 Electrochemical detection can

be used to achieve low detection limits in immunoassays with

little or no sample preparation, and atto- and zeptomole

detecting electrochemical immunoassays have been constructed.7,8

In homogeneous immunoassays, which have no separation

step to isolate the antibody–antigen complex from the unbound

assay constituents, electrochemical detection is not affected by

sample components such as chromophores, fluorophores, and

particles that often interfere with spectrophotometric detec-

tion. Therefore electrochemical measurements can be made on

colored or turbid samples such as whole blood, without

interference from fat globules, red blood cells, hemoglobin,

and bilirubin.9,10

Electrochemical techniques are generally organized into

three main categories of measurement: current, potential and

impedance. This article focuses primarily on those techniques

that measure current since they are the most commonly used in

biosensors.

1.2.1 Voltammetry/amperometry. Voltammetric and ampero-

metric techniques are characterized by applying a potential to a

working (or indicator) electrode versus a reference electrode and

measuring the current.11 The current is a result of electrolysis

by means of an electrochemical reduction or oxidation at the

working electrode. The electrolysis current is limited by the mass

transport rate of molecules to the electrode.11

The term voltammetry is used for those techniques in which

the potential is scanned over a set potential range. The current

response is usually a peak or a plateau that is proportional to

the concentration of analyte. Voltammetric methods include

linear sweep voltammetry, cyclic voltammetry, hydrodynamic

voltammetry, differential pulse voltammetry, square-wave

voltammetry, ac voltammetry, polarography, and stripping

voltammetry.11 These methods have a wide dynamic range,

and are useful for low level quantitation.

In amperometry, changes in current generated by the

electrochemical oxidation or reduction are monitored directly

with time while a constant potential is maintained at the

working electrode with respect to a reference electrode.5 It is

the absence of a scanning potential that distinguishes ampero-

metry from voltammetry. The technique is implemented by

stepping the potential directly to the desired value and then

measuring the current, or holding the potential at the desired

value and flowing samples across the electrode as in flow

injection analysis. Current is proportional to the concentration

of the electroactive species in the sample. Amperometric

biosensors have additional selectivity in that the oxidation or

reduction potential used for detection is characteristic of the

analyte species.1

Amperometric detection is commonly used with biocatalytic

and affinity sensors because of its simplicity and low LOD.12

Advantageously, the fixed potential during amperometric

detection results in a negligible charging current (the current

needed to apply the potential to the system), which minimizes

the background signal that adversely affects the limit of

detection. In addition, hydrodynamic amperometric techniques

can provide significantly enhanced mass transport to the

electrode surface,11,13 for example when the working electrode

moves with respect to the solution by rotating or vibrating,14,15

or in flow conditions where the sample solution passes over

the stationary electrodes.13,16,17 Electrochemical detection in

flow systems can be used in environmental monitoring and


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industrial processes more easily than steady-state batch

systems, since the flow conditions allow the solution to be

changed more easily in multistep assay procedures, and are

ideal for on-line monitoring.

Electrochemical sensors are part of an electrochemical cell

that consists of either three electrodes or two electrodes.

A typical three electrode electrochemical cell consists of a

working (or indicator) electrode of a chemically stable solid,

conductive material, such as platinum, gold, or carbon

(e.g. graphite); a reference electrode, usually consisting of

silver metal coated with a layer of silver chloride (Ag/AgCl);

and a platinum wire auxiliary electrode. The reference

electrode is usually further removed from the site of the redox

reaction in order to maintain a known and stable reference

potential.3 One advantage of this system is that the charge

from electrolysis passes through the auxiliary electrode instead

of the reference electrode, which protects the reference

electrode from changing its half-cell potential. A two electrode

system has only the working and reference electrodes. If the

current density is low enough (omA cm�2) then the reference

electrode can carry the charge with no adverse effect.5 Both

three electrode systems and two electrode systems are used for

sensors. However, two electrodes are generally preferred for

disposable sensors because long-term stability of the reference

is not needed and the cost is lower.

These electrodes can be easily miniaturized, so dimensions

on the order of micrometres are common, while nanometre

sizes have been demonstrated.18–20 Nanowires, nanoparticles,

and carbon nanotubes are now being incorporated into bio-

sensors. Shrinking electrode dimensions may lead to higher

sensitivity.3 Very small sample volumes (on the order of

microlitres and less) are required to detect with such small

electrodes due to their small surface areas, and this is a signifi-

cant advantage when the sample sizes are limited.21,22 Further-

more, electrochemical detectors and their required control

instrumentation can be easily miniaturized at a relatively low

cost by micromachining, making possible the manufacture of

field-portable instruments for biosensing. Since the limiting

current in voltammetry is temperature-dependent, the detec-

tion cell should be maintained at a constant temperature for

running calibrants and samples in order to obtain accurate

and precise results.23

Screen-printed electrodes (SPEs), patterned minielectrode

systems with working, reference and auxiliary electrodes, have

gained popularity in electrochemical biosensors due to their

low cost and ease and speed of mass production using thick

film technology.6 An SPE for detecting oxygen is shown in

Fig. 2. SPEs can also be miniaturized easily making them an

attractive transducer choice for microfluidic systems and

portable meters. The patterned working electrode is typically

made of conductive carbon ink that results in a rough surface

that makes difficult the exact determination of electrode

area.24 Gold coated and gold-based SPE sensors have been

used in stripping voltammetry to determine trace levels of lead,

copper, cadmium, and mercury in water samples.25 Nafion

coated SPE biosensors with immobilized butyrylcholinesterase

have also been developed to detect low levels of pesticides.26

Disposable SPEs have also been used in immunochemical

sensors and to measure blood glucose.27

Interdigitated array (IDA) electrodes are good amperometric

electrochemical transducers in biosensors (Fig. 3). IDAs are

made of two pairs of working electrodes consisting of parallel

strips of metal fingers that are interdigitated and separated by

insulating material.6,28 One electrode array serves as an anode

for oxidation and the other as a cathode for reduction as shown

in Fig. 3 for one anode finger and the adjacent cathode fingers.

The main advantage of using an IDA is the redox cycling of the

electroactive enzyme product or mediator that occurs when

different potentials are applied to the two electrodes causing

oxidation–reduction cycling when the electrode reaction is

reversible. The redox cycling provides lower limits of detection

because the current due to oxidation of each redox active

molecule contributes multiple times to the detection current.6,28

As a result, the signal-to-noise ratio is improved significantly

and a lower detection limit is obtained. Signal enhancement

increases as the spacing and width of the metal fingers decrease

because the diffusion distances for the redox species are shorter.

Typical signal enhancements provided by the IDA are about

3–10� and can be up to 1000� depending on the dimensions of

the IDA.28 IDA electrodes have been used as detectors in

electrochemical immunoassays.29 An IDA with dimensions on

the nanoscale was used for immunoassay detection of a virus.30

Fig. 2 Diagram of a screen printed electrode (SPE). Ref., reference

electrode; Aux., auxiliary electrode; and Work., working electrode.

Fig. 3 Cycling of a redox active species at the interdigitated array

electrode (IDA). Alkaline phosphatase (ALP) hydrolyzes o-phosphate

from a p-aminophenyl phosphate under alkaline conditions. R is the

reduced p-aminophenol (PAP). O is the oxidized p-quinone imine

(PQI).


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1.2.2 Impedance. Electrochemical impedance spectroscopy

(EIS), described by Lorenz and Schulze in 1975,31 measures the

resistive and capacitive properties of materials upon perturba-

tion of a system by a small amplitude sinusoidal ac excitation

signal typically of 2–10 mV.5,32 The frequency is varied over a

wide range to obtain the impedance spectrum. The in-phase and

out-of-phase current responses are then determined to obtain

the resistive and capacitive components of impedance, respec-

tively. Impedance methods are powerful because they are

capable of sampling electron transfer at high frequency and

mass transfer at low frequency. Impedimetric detection is

primarily used for affinity biosensors.27 It can be used to

monitor immunological binding events such as antibody

(Ab)–antigen (Ag) binding on an electrode surface, for example,

where the small changes in impedance are proportional to the

concentration of the measured species, the Ag.

The surface of the electrode can be modified by a highly

specific biological recognition element. In one approach the

recognition elements are incorporated in a conductive polymer

film formed on the surface of a working electrode by electro-

chemical deposition (Fig. 4). During the detection step, a

known voltage is applied to the electrode and the resulting

current is measured. The electron transfer resistance at the

interface between the electrode and the solution changes

slightly when analyte binds. Directly monitoring the formation

of an antibody–antigen conjugated layer provides a label-

free detection system with many potential advantages such

as higher signal-to-noise ratio, ease of detection, lower assay

cost, faster assays and shorter detector response times. How-

ever, regenerating the sensing surface for a subsequent measure-

ment in an impedance biosensor is typically very time-consuming

and not reproducible.27 This continues to be the biggest

limitation of immunosensors involving Ab–Ag complexes with

high affinity constants. The regeneration conditions can also

damage and release the immunoreagent bound to the surface

of the transducer.27

Electrochemical biosensors using impedance spectroscopy

to detect analytes have recently gained popularity among

the biosensor community.3 EIS has some advantages over

the widely used amperometric detection. The active site

participating in the biologically mediated redox reaction must

be easily accessible to the analyte solution and in close

proximity to the electrode surface. As discussed before, redox

mediators have been used to help overcome the accessibility

and proximity limitations but cause the detection to be limited

by the mediator’s mass transfer rate. Furthermore, some

additional redox active species such as urate and ascorbate

that are often present in the sample matrix can contribute to

the amperometric signal if the detection potential is not care-

fully chosen. Being directly able to impedimetrically monitor

the Ab–Ag binding helps by-pass the aforementioned limita-

tions. EIS is also insensitive to most environmental distur-

bances. However, biosensors using impedance detection

have to be carefully designed to minimize nonspecific binding

of the analyte. Nonspecific binding in affinity sensors will be

discussed further in Section 3.1.7. Using nanomaterials

such as gold nanoparticles and carbon nanotubes in electro-

chemical impedance sensors is advantageous due to the increased

electrode surface area, improved electrical conductivity of

the sensing interface, chemical accessibility to the analyte, and

electrocatalysis.32 Recent applications of impedance spectro-

scopy in affinity sensors will be described in Section 3.1.8.

1.2.3 Conductometry. Conductometric detection monitors

changes in the electrical conductivity of the sample solution, or

a medium such as nanowires, as the composition of the

solution/medium changes in the course of the chemical reac-

tion. Conductometric biosensors often include enzymes whose

charged products result in ionic strength changes, and thus

increased conductivity. Conductometry has been used as the

detection mode in biosensors for environmental monitoring

and clinical analysis. A conductometric tyrosinase biosensor

was developed to measure ppb amounts of pollutants such as

diuron, and atrazine and its metabolites.33 Conductometric

immunosensors have also been developed to detect foodborne

pathogens such as enterohemorrhagic Escherichia coli O157:H7

and Salmonella spp., which are of concern to biosecurity.34

The sensitive, low volume biosensor consists of an immuno-

sensor that is based on an electrochemical sandwich immuno-

assay, and a reader device for measuring the signal.34 Drug

detection of methamphetamine in human urine has also been

done using conductometry.35

1.2.4 Potentiometry. Potentiometric sensors are based on

measuring the potential of an electrochemical cell while drawing

negligible current. Common examples are the glass pH elec-

trode and ion selective electrodes for ions such as K+, Ca2+,

Na+, Cl�.1,2 The sensors use an electrochemical cell with

two reference electrodes to measure the potential across a

membrane that selectively reacts with the charged ion of

interest. These chemical sensors can be turned into biosensors

by coating them with a biological element such as an enzyme

that catalyzes a reaction that forms the ion that the underlying

electrode is designed to sense. For example, a sensor for

penicillin can be made by coating a pH electrode with

penicillinase, which catalyzes a reaction of penicillin that also

generates H+.36 The pH electrode senses the change in pH at

its surface, which is an indirect measure of penicillin.Fig. 4 A diagram of an Ab–Ag affinity sensor with impedimetric

detection.


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Field effect transistors have been adapted to chemical

sensors (ChemFETs) by incorporation into an electrochemical

cell.37,38 They can also be made into biosensors by coating the

sensing surface with a biological agent such as described above

for penicillin.39 The light addressable potentiometric sensor

(LAPS) determines the surface potential optically by means of

the photovoltaic effect.40 The LAPS can also be used as a

biosensor by adding a biological element to its surface, such as

an oligonucleotide.41

1.2.5 Miniaturized electrochemical transducers. Miniaturi-

zation is a growing trend in analytical chemistry. In order to

design and manufacture small biosensors, the transducer or

the electrode needs to be small and portable. The manufacturing

capabilities for depositing microelectrodes on surfaces are

good and microelectrodes can easily be deposited on a micro-

fluidic chip or other solid surface using vapor deposition.6

Usually the electrode is part of a bigger device such as a

handheld meter or a microfluidic system.

Microelectrodes are defined as electrodes with a diameter in

the micrometre scale, and are made as disks or cylinders from

carbon fibers or metal microwires.18,19 The first measurements

using microelectrodes to measure the concentration of oxygen

in biological tissues were made in early 1940s,42 and they have

since been used to measure electroactive species in critical

places such as inside a mammalian brain.18 Measurements

with voltammetric microelectrodes have been made even inside

a very small, live biological cell.43 This is because the important

reactions occur at the microelectrode surface instead of bulk

solution, and the very small sensing surface area of a micro-

electrode can be easily inserted into very small drops or spaces

without causing much disturbance or damage. Carbon fiber

microelectrodes have been used to detect 190 zmol of catechol-

amine release from a single, stimulated rat nerve cell,44 to

directly monitor catecholamines released from adrenal cells in

culture,45 and to measure the release of serotonin from neuronal

vesicles achieving a 4.8 zmol detection limit.46 Microelectrodes

have also been used as detectors in microvolume electro-

chemical immunoassays.22 The nanoamp to picoamp currents

generated at microelectrodes are so small that they are virtually

nondestructive,18 and amplification of the small currents produced

is typically required in order to observe the signals.6

2. Biocatalytic sensors

2.1 Introduction to enzyme-based electrodes

Enzyme electrodes are electrochemical probes with a thin

layer of immobilized enzyme on the surface of the working

electrode.47,48 The enzyme is the most critical component of

the enzyme electrode since it provides the selectivity for the

sensor and catalyzes the formation of the electroactive product

for detection.49 The electroactive product can be monitored

directly using amperometry, in which the produced current is

measured in response to an applied, constant voltage. Alter-

natively, the disappearance of the redox active reactant in an

enzyme-catalyzed reaction can be monitored by the electrode.

The activity of the immobilized enzyme depends on solution

parameters and electrode design. The rapid enzymatic

catalysis can also sometimes provide significant signal

amplification in a biosensor.5 The shelf life and stability of

an enzyme generally determine the lifespan of the biosensor.

The use of enzyme electrodes as biosensors will continue to

increase because they are simple and inexpensive to construct,

they provide rapid analysis, they easily regenerate, and they

are reusable.2,5 However, the number of available enzyme-

based biosensors is still smaller than the number of potential

analytes. Another disadvantage of enzyme electrodes is that

the enzyme layer in the biosensor has to be replaced periodi-

cally since it gradually loses activity. Also, clever electro-

chemical detection strategies or membranes are sometimes

required to prevent interference from other redox active

species at certain detection potentials.

Development of biocatalytic sensors for medical appli-

cations, primarily blood glucose monitoring starting in the

late 1960s, was the main driving force for this research area.5

Enzyme-based biosensors can be historically divided into three

generations. First-generation biosensors were oxygen-based

whereas second-generation are mediator-based. Third-generation

biosensors are so-called directly coupled enzyme electrodes.

Electrodes coated with glucose oxidase (GOx) have been

widely used in detection of glucose since the pioneering work

of Clark and Lyons in the 1950s and 1960s (Fig. 5).50 These

amperometric sensors became known as the first-generation

biosensors or Clark oxygen electrodes and were soon imple-

mented by Updike and Hicks, who constructed the first func-

tional biocatalytic sensor for glucose.51 In the first-generation

biosensors, an oxidase enzyme is immobilized behind a semi-

permeable membrane at the surface of a Pt electrode.

GOx is a readily available, inexpensive, and stable enzyme

from Aspergillis niger that is among the most important

enzymes in biosensor applications and industrial processes.

GOx is highly specific for b-D-glucose, which can be detected

via the following reactions.2,5,52

b-D-Glucose + GOx–FAD - GOx–FADH2

+ d-D-gluconolactone (1)

GOx–FADH2 + O2 - GOx–FAD + H2O2 (2)

H2O2 - 2e� + O2 + 2H+ (3)

Fig. 5 Oxygen-dependent first-generation biosensor with ampero-

metric detection.


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In eqn (1) the prosthetic group of the enzyme, FAD, is reduced

and glucose is oxidized to d-D-gluconolactone. Molecular

oxygen acts as the oxidizing agent to produce hydrogen

peroxide (eqn (2)). During the oxidation of H2O2 at a working

electrode two electrons are transferred directly to the electrode

(eqn (3)), resulting in the current response of the enzyme

electrode. These first-generation sensors required the ample

and constant presence of ambient oxygen as a co-substrate for

the enzyme to function optimally. However, oxygen is not very

soluble in aqueous solutions and can therefore limit the

currents produced in the presence of the analyte.

Direct redox reactions between enzymes and electrodes are

very rare because most proteins tend to denature at the

electrode surface and many direct electron transfer reactions

are slow and irreversible.1 However, a limited number of

enzymes such as horseradish peroxidase have proven capable

of direct electron transfer between the enzyme active site’s

prosthetic group and the electrode.53 The active site of an

enzyme that allows the selective targeting of an analyte is

usually buried within the enzyme’s tertiary protein structure,

near the centroid of the protein.4 Therefore, the electrons

produced in the enzyme-catalyzed reaction cannot always be

easily and rapidly transferred to the electrode surface thereby

limiting the electrical communication between the enzyme and

the transducer. The widely accepted Marcus theory of electron

transfer states that electron transfer decays exponentially with

distance.54,55 Therefore enzymes often require some assistance

with electron transfer to the transducer surface.

Artificial redox mediators are small, soluble molecules

capable of undergoing rapid and reversible redox reactions,

which shuttle electrons between the redox center at the active

site of the enzyme and the electrode surface. Mediators have

replaced O2 molecules as the electron shuttle (eqn (4)) in

glucose sensors. Mediators are re-oxidized at relatively low

potentials and generate a current when they come in contact

with the working electrode (eqn (5)).

GOx–FADH2 + 2MediatorOx - GOx–FAD

+ 2MediatorRed + 2H+ (4)

2MediatorRed - 2MediatorOx + 2e� (5)

Mediators should ideally be nontoxic, independent of the

pH, stable in both the oxidized and reduced forms, and

unreactive with oxygen.1 Although many organic compounds

are capable of acting as enzyme mediators, organometallic

redox compounds are the most common.1,2 Examples of

previously used mediators include quinones, organic conducting

salts, dyes, ruthenium complexes, ferrocene, and ferricyanide

derivatives. Mediated enzyme electrodes had a much better

sensor performance than the first-generation biosensors

mainly due to eliminating the O2 dependence and being able

to control the concentration of the oxidizing agent in the

biosensor.1 Hand selecting the oxidizing agent for the sensor

also allowed more suitable oxidation potentials to be used for

the amperometric sensors. These mediated enzyme electrodes

were named second-generation biosensors.

By carefully selecting a mediator and a suitable redox

potential, the transduction event at the second-generation

biosensor could be measured in a potential range where other

possible sample components such as ascorbate, urate, and

paracetamol are not oxidized or reduced thereby minimizing

interferences.5 Incorporating redox mediators also allowed

other oxidoreductase enzymes such as peroxidases and dehydro-

genases to be used as the biorecognition element in the sensor

thereby expanding the list of possible target analytes.

Third-generation biosensors have the biorecognition com-

ponent coupled with the electrode by co-immobilizing the

enzyme and the mediator at an electrode surface. This can

be achieved by direct electrical contact between the enzyme

and the electrode, immobilizing the enzyme and mediator in a

conducting polymer, or ‘wiring’ the enzyme to the electrode by

immobilizing it in a redox polymer (Fig. 6) as first described by

Heller et al.56,57 The co-immobilization prevents the mediators

from diffusing out of the biosensor film. The co-immobilized

mediators, or the flexible surrounding redox polymer, help to

transport electrons between the enzyme’s active site and the

working electrode surface in an array of rapid electron relays

and hence generate high current densities.2 The enzymes

immobilized in flexible redox polymers that are covalently

attached to the electrode have been called ‘wired enzymes’.

The 3rd-generation sensors are ideal for repeated measure-

ments since neither mediator nor enzyme need to be added.

This self-contained nature also lowers the cost per measure-

ment and opens up possibilities for continuously monitoring

the analytes.

2.2 Preparing enzyme electrodes

2.2.1 Methods for immobilizing enzymes to electrode surfaces.

Enzyme electrodes have been studied extensively and various

physical and chemical schemes have been used to immobilize

enzymes on the electrochemical transducer. The objective is to

have an intimate contact between the enzyme and the trans-

ducer’s sensing surface without blocking the active site of the

enzyme or drastically altering the enzyme geometry.2 Immobili-

zation methods are considered successful if the biosensors pre-

pared are stable, reusable, and maintain the selectivity of the

enzyme. Although immobilization may alter the conformation of

the enzyme, thereby reducing its activity, many methods have

been successful. Some immobilization methods even improve

enzyme stability by minimizing enzyme unfolding. The enzyme

should have high Vmax and low Km values when immobilized on

Fig. 6 Third-generation catalytic biosensor containing enzymes

wired to the electrode through a conducting redox polymer.


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the transducer.2 Vmax is the maximal velocity of a reaction that

occurs at high substrate concentrations when the enzymes are

saturated. By having an immobilized enzyme with a high Vmax

the electrochemical transducer responding to a reaction catalyzed

by the enzyme has a broader range where the signal is propor-

tional to substrate concentration for reliable quantitation of the

analyte. Km, the Michaelis constant, is the substrate concen-

tration at which the reaction velocity is half-maximal. Enzymes

with low Km reach maximal catalytic efficiency at low substrate

concentrations. The immediate environment around the immobi-

lized enzyme can be carefully designed to enhance the enzyme

activity and the overall biosensor performance.

The easiest approach is to physically entrap a solution of

the enzyme between preformed membranes on the electrode

surface.2 The inner membrane protects the electrode surface

from interfering substances and electrode fouling due to

adsorption. The outer membrane also provides some selecti-

vity based on the pore size or chemical nature of the polymer,

stabilizes the sensor response by moderating the substrate

diffusion to the enzyme layer, and provides a biocompatible

outer surface for the sensor.5 In physical immobilization

methods the native composition of the enzyme is preserved

since the methods do not involve the formation of covalent

bonds.34 Chemical methods involve the formation of covalent

bonds between the functional groups of the enzyme and the

electrode material.5 Common enzyme immobilization methods

include enzyme entrapment against the electrode using a

preformed membrane; encapsulation; inclusion in a gel or

electropolymerized film; incorporation in a carbon paste;

and using biospecific interactions such as biotin–avidin binding,

adsorption, cross-linking, and covalent attachment (Fig. 7).58,59

Covalent bonding provides the most stable immobilization of

proteins followed by cross-linking and encapsulation.1 Covalent

bonding to the transducer links functional groups on the

enzyme such as NH2, COOH, OH, and SH that are not

necessary for the catalytic activity of the enzyme. The coupling

reactions need to be done under mild conditions (low ionic

strengths, low temperatures, and near physiological pHs) and

often in the presence of the enzyme–substrate in order to protect

the catalytic activity of the enzyme.1 Adsorption is the least

stable of the common immobilization methods.1 The forces

linking the biorecognition element to the transducer in adsorp-

tion are primarily very weak van der Waals forces with occa-

sional hydrogen bonds that are not very stable or permanent.1

Therefore the lifetime of a sensor prepared using adsorption is

rather limited. However, adsorption is very easy because it does

not require any reagents or clean-up and is less disruptive to the

enzymes.1 The formation of intermolecular interactions with

the surface may compete with similar interactions stabilizing the

enzyme, and is often a prelude to denaturation. This is probably

why adsorption usually works best in the short term, because

the protein deformation increases with time. Adsorption is

often sufficient for short-term studies. The stability of immobi-

lized enzymes with respect to time, temperature, and pH is

typically greater making enzyme electrodes preferable to soluble

enzyme assays.5,52 Covering the immobilized enzyme layer with

a membrane or a polymer coating also helps to minimize

interferences by physically blocking some interfering species

from approaching the electrode surface.52

2.2.2 Optimizing enzyme electrodes. Although many enzyme

electrodes have been fabricated and some sensors have reached

the commercial stage, some factors that prevent their wider

adaptation and successful routine use still remain. Research

continues in trying to overcome the dependence of enzyme

activity on the solution conditions such as temperature, pH,

ionic strength, and buffer composition.60 Ideally the solution

conditions should remain constant between samples and during

the measurements. The enzyme electrode should also have a

wide linear range. Enzymes become saturated with their sub-

strate at high concentrations due to their active sites becoming

the limiting reagent, thus causing the signal response to no

longer be proportional to the analyte concentration. The

amount of enzyme incorporated into the sensor can however

be adjusted based on the expected sensor application. The

catalytic biosensor should also be biocompatible since blood

and other biological fluids are the most common sample

matrices for enzyme electrodes. Many blood components foul

the electrode in a matter of minutes unless special precautions

are taken in designing the sensor’s outermost surface properties

and permeability to prevent the adsorption of sample com-

ponents on the electrode surface.60 Product design requirements

also include optimization of sample introduction, sample size,

the sensor’s reproducibility, selectivity, sensitivity, stability, cost,

and ease of use.5 The storage stability of enzymes immobilized

on electrode surface varies from hours to months depending on

the sensor preparation and design, and the storage environment.

2.3 Examples of biocatalytic sensors

2.3.1 Glucose sensors. Enzyme electrodes are produced

commercially and are routinely used in biomedical appli-

cations such as glucose testing in clinical laboratories and

personal monitoring by diabetic patients.2,5,48 Low cost blood

glucose home monitoring kits consisting of handheld battery

operated meters and disposable glucose test strips based on

glucose oxidase (GOx) or glucose dehydrogenase enzyme

electrodes are sold off the shelf worldwide. Biosensors for this

application must be easy to use, reliable, and inexpensive.1 In a

typical sensor, a single drop of blood is placed on a disposable

PVC sample strip on which the dry reagents have been

deposited using a method similar to ink-jet printing techno-

logy. The test strip also contains two electrodes, one holding

the enzyme and a mediator for the amperometric detectionFig. 7 Common methods of immobilizing enzymes onto an electrode

surface.


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of glucose, and the other serving as a reference electrode.

The current produced when a potential is applied gives a read-

out to a liquid crystal display on the glucose meter. The

commercially sold glucose test strips are second- or third-

generation biosensors and no longer rely on oxygen as the

oxidizing agent. Ferricyanide is a commonly used mediator for

the second-generation sensors. Eqn (1), (4), and (5) describe

the sequence of reactions for such a sensor when GOx is used

as the enzyme. The commercially sold blood glucose meters

typically have a range of 1.1–33.3 mM glucose with a precision

of �3–8% and test time of about 30 seconds or less.1

Some invasive and minimally invasive implantable glucose

sensors that have an intimate contact between the biological

fluids or tissues and the biocatalytic sensor have been developed.5,61

The minimally invasive blood glucose sensors are inserted

subcutaneously into the arm or belly of a patient. More

invasive, intravascular sensors that measure glucose levels in

hospitalized diabetes patients are also being developed. Some

problems such as pain, skin irritation, limited lifetime of the

sensor, and accuracy of the data continue to slow down their

wider use.62

2.3.2 Xanthine sensors. Xanthine oxidase (XO) catalyzes the

oxidation of xanthine to uric acid (eqn (6)). Amperometric

biosensors using immobilized XO are highly specific for xanthine,

which can be measured by the following redox reaction:

Xanthine + O2 + XO - uric acid + H2O2 + XO (6)

Xanthine is an intermediate of purine metabolism and is

produced after adenosine triphosphate (ATP) decomposition.

The physiological conversion of xanthine by xanthine oxidase

is of increasing medical interest.63 Moreover, xanthine sensors

are frequently used in food industries to determine the freshness

of fish. The need for maintaining an acceptable quality of

fish sold to consumers requires rapid and reliable analytical

methods that detect the products formed in their degradation

processes. After the death of a fish, nucleotides such as ATP

are most affected by degradation and give rise to the formation

of inosine, which is transferred to hypoxanthine by action of

the enzyme nucleoside phosphorylase.64 Hypoxanthine causes

a bitter taste in the degrading meat.65 XO catalyzed oxidation

of hypoxanthine to xanthine and conversion of xanthine to

uric acid occurs in two steps.64 The quantitation of xanthine or

hypoxanthine can therefore be used to determine the freshness

of fish.64,65 Other existing methods for detecting xanthine or

hypoxanthine such as anion-exchange chromatography, thin

layer chromatography, precipitation and capillary electro-

phoresis are complicated and very time-consuming. Therefore

biocatalytic sensors with amperometric detection continue to

be developed to monitor the freshness of fish meat.66–68

2.3.3 Lactate sensors. Lactate, an ester of lactic acid, is a

product of fermentation and is produced during cellular

respiration as glucose is broken down. Its concentration in

blood rises from the normal value of 0.9 mM to about 12 mM

due to strenuous exercise such as running, which results in

anaerobic metabolism.69 Small handheld electrochemical

lactate meters for use in sports medicine capable of inter-

mittent ‘‘spot’’ lactate monitoring are being manufactured by

Senslab (Germany) and Arkray (Japan).70 These sensors

require only 0.5 mL and 5 mL blood samples, respectively.

The concentration of lactate in blood is also a sensitive

measure of oxygen deprivation from ischemia, trauma, and

hemorrhage, which can lead to life-threatening shock, and its

measurement has therefore become a vital component in

medical monitoring.70 Blood lactate levels are used as indi-

cators of conditions such as acidosis or bacterial meningitis.71

Conventional photometric assays for lactate are slow and not

suited for continuous lactate monitoring systems that are

being developed for medical applications. Bench top lactate

biosensors are also routinely used to measure lactic acid in

milk and other foods.

Four different enzymes have been used as the biorecognition

component in lactate biosensors: lactate dehydrogenase, lactate

oxidase, lactate monooxidase, and cytochrome b2.1 Some of the

electrochemical lactate sensors include mediators such as

NAD+/NADH and ferricyanide.1 All of these enzymes help

ultimately to produce a current at the working electrode that is

measured amperometrically.

2.4 Interference-based enzyme electrodes

Interference-based enzyme electrodes are probes used for quanti-

tative analysis based on the changes in the rate of catalytic

reactions when enzyme effectors bind, such as inhibitors or

activators.72 Usually, the binding of an inhibitor to the enzyme’s

active site or another site that causes the conformation of the

enzyme to be altered, results in a decreased electrochemical

response because substrate cannot freely access the catalytic site.

Sensors using this operating principle have been developed for

pesticides such as organophosphate and carbamate, respiratory

poisons such as cyanide and azide, as well as toxic heavy metals.2

Enzymes that have been used in these enzyme electrodes include

tyrosinase, horseradish peroxidase, and acetylcholinesterase. Due

to the operating mechanism and the uses these sensors have, they

have also been called enzyme inhibition biosensors or toxin

biosensors.2,72

2.5 Biosensors based on tissue and bacteria

Some biocatalytic sensors incorporate cellular materials such

as plant tissues as the recognition component.2 These bio-

catalytic electrodes function in a manner similar to that for

conventional enzyme electrodes (i.e., enzymes present in the

tissue or cell produce or consume electrochemically detectable

species). Whole cells and tissue slices are sometimes a better

source of enzymatic activity compared to isolated enzymes as

some enzymes are expensive or not commercially available in

the pure state.2 Also, many isolated enzymes have limited

stability and lifetime compared to enzymes in their native

environment. However, the sensor response may be slower for

these sensors because there is more tissue material for the

substrate to diffuse through.1

Bananatrode, a banana tissue containing electrode, was one

of the early uses of tissue in a biosensor.63 The banana tissue,

which is rich with polyphenol oxidase (PPO), can be mixed in a

carbon paste matrix to yield a fast responding and sensitive

dopamine sensor. The amperometric probe has high biocatalytic

activity, good time stability, and favorable selectivity.73


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Live microorganisms have also been coupled with electro-

chemical transducers (i.e. electrodes) to monitor biotechno-

logical processes such as brewing, food manufacturing,

waste-water treatment, energy production, and pharma-

ceutical synthesis.1,2,74 Bacteria are often immobilized on

transducers by microencapsulation where an inert membrane

is used to trap the microbe on the electrode surface.1 Changes

in the respiration activity of the microorganism, induced by

the target analyte, results in a lower surface concentration of

electroactive metabolites (e.g., oxygen), which can be detected

by the electrochemical transducer.2,74 Some microorganisms

also produce electroactive metabolites that can be monitored

directly.1 Using microbes in biosensors gained popularity

because they are typically cheaper to obtain than isolated

enzymes, are less sensitive to inhibition by other sample com-

ponents, and are more tolerant of slight temperature and pH

variations than enzymes.1,5 Some of their disadvantages include

longer recovery times after exposure to the analyte of interest,

longer response times, hysteresis effects, and possible loss in

selectivity due to containing many types of enzymes.1,5,74

3. Affinity biosensors

3.1 Immunoassays and immunosensors

Immunoassays gained popularity for biomedical applications

in the 1970s because of the impressively low detection limits

and high selectivities for analyzing complex samples that could

be achieved with relatively simple procedures and instru-

mentation. The availability of highly selective antibodies for

an increasingly wide variety of important analytes was also an

important factor in the growth of the method over the

following decades. The development of more sensitive labels

and detection devices also improved the sensitivity of the

assays even further. Once immunoassays became more com-

mon, the development of more convenient immunosensors

that are easier and faster to use gained momentum.

Most applications of immunoassays (IA) and immuno-

sensors with electrochemical detection were initially developed

at research laboratories due to the level of expertise required,

time, and the high initial cost of developing and optimizing a

new immunoassay.27 However, the cost of immunological

reagents continues to decrease with recent developments in

molecular biology techniques. Many of the early radioimmuno-

assays were developed for biomedical applications such as

detecting hormones and disease related proteins, but appli-

cations in environmental, agricultural, processed food and

beverage areas, and to detect harmful chemical and biological

agents in national defense, have become more common.6,27

The advantages of IAs such as exceptionally high specificity

of Ab for Ag, small sample volumes, low detection limits, little

or no sample preparation, reduced use of chemicals, little

waste, and ease of automation, far outweigh their limitations,

thus making the IAs an attractive alternative to the more

conventional quantitative analytical methods like chromato-

graphy and mass spectrometry.6 Many IA formats also allow

the simultaneous analysis of multiple samples, which improves

efficiency and makes the assays relatively fast and cost effec-

tive. The immunoassays and affinity biosensors are relatively

easy to use once fully developed and optimized, and the

reductions in chemicals used, waste disposal, expensive instru-

ments and maintenance also help lower the overall cost per

analysis. The four key factors involved in the design of a

sensitive immunoassay are its format, the type of label, the

method of detection, and being able to minimize nonspecific

binding (NSB).27 These factors will be discussed further in

Sections 3.1.4–3.1.7.

3.1.1 Biorecognition and immunochemical reactions. IgG

antibodies (Ab), large Y-shaped glycoproteins of MW E150 kDa, are produced by a host in response to the presence

of a foreign molecule called antigen (Ag).5 Antigens are any-

thing that the body recognizes as foreign such as chemical

compounds, proteins, and particulate matter (dust, pollen,

etc.).75 Abs are produced by specialized B lymphocyte cells

of the immune system and can usually be found in blood

serum, tissue fluids, and membranes of vertebrates.75 Antigens

commonly have relatively high molecular weights, are recog-

nized as nonself or foreign by the immune system and have a

certain level of chemical complexity.75 For example, synthetic

homopolymers composed of a single sugar or amino acid tend

to lack immunogenicity regardless of their large size due to a

lack of structural complexity.75 The production of Abs against

low molecular weight analytes (MW o 1000 g mol�1) called

haptens is more challenging and often requires coupling the

hapten to a carrier protein with a spacer molecule before an

immune response can be provoked in the host animal.75,76

IgGs have four polypeptide chains (two identical heavy

chains with MW of 50 000 or higher and two identical and

smaller light chains with MW of about 25 000) that are held

together by disulfide bonds and noncovalent interactions such

as hydrogen bonds as seen in Fig. 8.75 Each chain has several

different domains. Each Ab molecule has two identical binding

sites and is therefore called bivalent. The highly selective

antigen-binding site is formed at the tips of each of the Y

arms where a heavy-chain variable domain (VH) and a light-

chain variable domain (VL) come close together.75 These

complementarity-determining regions (CDRs) are the domains

that differ most in their sequence and structure between

different antibodies. Parts of VL and VH contribute to the

finger like loops that interact with the antigens.75 The non-

covalent interaction between the Ab and Ag is highly specific,

which makes antibodies an excellent biorecognition element

for affinity biosensors. However, unlike in enzyme–substrate

interactions, Ab–Ag binding does not lead to an irreversible

chemical alteration in either the Ab or the Ag.75 The non-

covalent interactions that are cumulative and form the basis

for the binding interaction include hydrogen bonding, ionic

bonds, hydrophobic interactions, and van der Waals forces. A

very close fit resulting from a high degree of complementarity

between the Ab and the Ag is required for the noncovalent

interactions to form since they operate over very short dis-

tances. Sometimes the exceptional selectivity can be a dis-

advantage when an Ab is selective only for one isomer of the

Ag when a sensor should ideally measure the total amount of

the Ag type.1

The unique antibody-binding region of the CDR is also

called the paratope, and recognizes and binds with high


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affinity to a complementary site on the antigen called the

epitope. The paratope–epitope complementarity is based on

size (in nm scale), shape, and the chemical compatibility within

the interface. The binding of Ab to an Ag is also very powerful

and affinity constants of about 106 are common for Ab–Ag

complexes, with some being considerably higher.1 In some

cases binding of Ag induces conformational changes in the Ab,

Ag, or both. This conformational change results in a closer fit

between the epitope and the antibody’s binding site, but may

incur an energetic cost thereby reducing the binding affinity.75

3.1.2 Antibody production. The production of the anti-

bodies (Abs) against a specific antigen (Ag) can be fairly

difficult and time-consuming.77 A small host animal such as

a mouse, a rabbit, or a chicken is injected with small sub-lethal

doses of Ag to challenge their humoral immune system to

produce the specific Abs against the foreign invader. Some-

times larger mammals such as goats are preferred as the host

because the amount of blood serum that can be collected is

greater. Mice are usually used in the initial stages of mono-

clonal Ab (MAb) production.

MAbs are produced by a single Ab-producing cultured cell

line (containing clones of a single parent cell) in a bioreactor

and are identical in the primary structure.75 MAbs can also be

produced in microbial systems and transgenic mice.78,79 These

homogeneous Abs that are known for their high specificity

and affinity are used as the primary or capture Ab in most

research, diagnostic, and sensing applications. MAbs have an

inherent specificity toward a single epitope that allows fine

detection and quantitation of small differences in Ag. Poly-

clonal Abs are a heterogeneous mixture of immunoglobulin

molecules secreted against a specific antigen, each recognizing

a different epitope.75 They have varying affinities for the Ag

and are often used as the secondary Ab in immunoassays.

The small size of mice prevents their use for sufficient

quantities of polyclonal, serum antibodies.77 Animals usually

used for polyclonal Ab production include chickens, goats,

rats, guinea pigs, hamsters, sheep, camels, llamas, and horses.

Rabbit is by far the most commonly used laboratory animal

for Ab production. The soluble antibodies produced by the

host in these immune system challenges are then recovered

from the harvested blood serum by a series of extractions and

purifications.75

3.1.3 Cross-reactivity. Although most Ab–Ag interactions

are highly specific, some antibodies produced against one Ag

can cross-react with an unrelated Ag.75 This type of cross-

reactivity occurs if two different antigens share identical or

very similar epitopes. Usually the antibody’s affinity for the

similar epitope is less than for the original epitope but the

cross-reactivity can still result in false positives and inter-

ference for the affinity sensor.

3.1.4 Immobilization of antibodies. Like enzymes, DNA,

and other biorecognition molecules, antibodies are very sensitive

to their environmental conditions.3 Typically Abs have to be

immobilized on a solid surface in a biosensor application

which can lead to loss of their biological binding activity.

Therefore, special care has to be taken when immobilizing

antibodies with respect to their orientation on the solid

surface. The tips of the Y-shaped arms containing the binding

sites of antibodies have to be exposed to the sample and

therefore Abs cannot be randomly oriented on the surface.

Nonspecific interactions between the surface and the mis-

oriented antibody can also lead to denaturation of the binding

sites. Also, the density of the Abs on the surface cannot be too

high to minimize steric hindrance.3,80 Common Ab immobili-

zation methods include biotin–streptavidin linkages,6,22,29 adsorp-

tion to a conductive polymer matrix such as polypyrrole,81 and

covalent binding.3,82

3.1.5 Formats for enzyme immunoassays. Enzyme immuno-

assays (EIAs) were first introduced by Engvall, Perlmann,

Van Weemen, and Schuurs in 1971 as an alternative to

radioimmunoassays.27 The previously used radioactive label

indicating that an Ab–Ag complex had formed was replaced

by a safer, selective and less expensive enzyme label at the cost

of less sensitivity and more complexity.27 In EIAs the activity

of the enzyme label in generating electroactive product is

measured. Enzymes are also highly selective for their given

substrate, and can provide a large signal amplification due to a

high turnover rate, which yields low limits of detection. How-

ever, as discussed in Section 1.1.1 the activity of the enzyme

labels can be affected by reaction conditions that have to be

controlled during the detection step. Like radioimmunoassays,

enzyme immunoassays can be time-consuming due to including

multiple incubation and washing steps. Many variations of

immunoassays have been developed that allow sensitive quanti-

tation of either Ag or Ab. The two main immunoassay (IA)

formats are homogeneous and heterogeneous.27 Homogeneous

assays, which do not contain separation steps, are faster and

easier, but have poorer limits of detection. Homogeneous assays

are also more susceptible to interferences by other species in the

sample than IAs with other formats.27 Heterogeneous assays

include a physical separation step to isolate the antibody–

antigen complex from the unbound constituents followed by a

wash step to remove any unbound materials. The separation

step in a heterogeneous assay makes the procedure longer, but

results in significantly better limits of detection.

Homogeneous and heterogeneous EIAs can be done either

competitively or noncompetitively.27 Competitive immunoassays,

Fig. 8 Y-shaped antibody structure. Ag, antigen; VH, variable region

of heavy chain; VL, variable region of light chain; CH1–3, constant

regions of heavy chain; and CL, constant region of light chain.


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also known as limited reagent assays, are often used when the

antigen is small and has only one epitope.75 In a competitive

assay a limited amount of Ab is used, which is insufficient to

bind with all the Ag molecules in the sample. A fixed, known

amount of labeled Ag is mixed with the unknown sample

and allowed to incubate. Unlabeled Ag and the labeled Ag

compete for binding to the limited number of capture Ab sites.

Rinses are required to separate the unbound Ag from the

bound prior to the detection step. A decrease in signal

response indicates the presence of the Ag in the sample

being analyzed. The ratio of limited Ab reagent to the added

labeled Ag must remain constant between samples to obtain

quantitative results.

Noncompetitive assays are also called excess reagent assays

and are better suited for large analyte molecules with several

epitopes.75 The Ag sample is incubated with an excess of Ab

reagent. All the Ag molecules form a complex with antibodies,

but not all of the Ab-binding sites are occupied. To detect the

amount of Ag attached to an Ab, a labeled secondary Ab

is added which binds to another, available epitope on the

bound Ag. This leads to the formation of a sandwich complex

(Ab:Ag:Ab*). Unbound excess reagent is washed away after

each incubation step. The electrochemical signal produced

during the detection step is directly proportional to the

amount of Ag in the unknown sample.

Sandwich IAs are often referred to as enzyme-linked immuno-

sorbent assays (ELISA) because the antibody or the antigen is

immobilized on a solid surface such as a bead, membrane, a

polystyrene well, or an electrode surface. Fig. 9 shows the main

steps in a sandwich enzyme immunoassay. Having the immuno-

reactants of the ELISA immobilized makes it easy to separate

bound from unbound material during the assay washing steps.27

3.1.6 Enzyme labels and substrates. The enzyme label

chosen for the IA with electrochemical detection should have

a high catalytic activity for the corresponding substrate and be

fairly stable in the sample matrix. It should also be readily

available in a purified and soluble form at a reasonable cost.

The enzyme label should contain surface functional groups

that can be used to form conjugates with other molecules as

needed without impairing its catalytic activity or compromising

the biorecognition events. The redox active product that is

formed by the enzyme catalysis should have a low redox

potential to minimize interference from other components in

the sample, while the substrate should be electroinactive at the

measuring potential to keep the background signal low.27 It is

usually not necessary to remove oxygen from the sample if the

observed reaction is an oxidation occurring between +200

and +900 mV. The lower end of the range is more desirable

because the more positive values may result in electrolysis of

the solvent. Several enzymes satisfy the above requirements

and are used in electrochemical IAs and immunosensors. The

most commonly used enzyme labels are alkaline phosphatase

(ALP), b-galactosidase (b-Gal), horseradish peroxidase

(HRP), and glucose oxidase (GOx).1,2,27 GOx has a lower

activity than the other enzyme labels and is typically used in

amperometric immunoassays where the product is detected

directly.

3.1.7 Nonspecific binding. Nonspecific binding (NSB)

involves the adsorption of conjugated enzyme or other labels

used for immunoassay to materials other than the analyte.27

This phenomenon, which increases the background signal, is

the major determinant of the detection limit of the IA and

therefore including procedures that minimize NSB in immuno-

assays is critical. NSB can be reduced with blockers such as a

nonionic surfactant, Tween 20, protein blockers such as

bovine serum albumin (BSA), polyethylene glycol,83 gelatin,84

casein,85 and proprietary blended commercial products. Self-

assembling monolayers of oligo(ethylene glycol)86–88 and

dextran layers89 have also been used successfully to prevent

NSB on affinity biosensor surfaces. These NSB blocker re-

agents are commercially available and widely used in affinity

biosensors.

With plastic surfaces, such as polystyrene used to make

beads and microtiter wells, hydrophobic interactions usually

dominate the adsorption process.8 The adsorption is entropi-

cally driven and can usually be minimized by physically coating

the exposed areas of the reaction vessel by surface treatments

such as a mixture of bovine serum albumin and a detergent

such as Tween 20.90 Sulfonate ion-pairing reagents have been

found to reduce NSB on positively charged surfaces.8 Deter-

gents and proteins can be added to the buffer to block NSB

with bead-based immunoassays.14,15 A 13-fold reduction

in detection limit has been seen in blocked electrochemical

immunoassays compared to the unblocked assays.90 Contact

between NSB blocking agents and the electrode transducer

should be avoided because the blockers may adsorb on the

electrode surface, fouling it.27

3.1.8. Applications of immunoassays. Immunoassays and

enzyme sensors have been incorporated into portable instru-

ments capable of quickly measuring multiple analytes. A good

example is the i-STATt, which is able to make measurements

on small volumes (17–95 mL) of whole blood.91 The i-STATt

analyzer is based on single-use disposable cartridges con-

taining a microfabricated biosensor array. The system auto-

matically calibrates the sensors and analyzes the sample.

Ion-selective electrodes are used to determine Na+, K+,

Cl�, Ca2+, pH and pCO2. Amperometric enzyme biosensors

are used to determine glucose, lactate and creatinine using the

principles described above. Recently, cartridges capable of

sandwich immunoassay with electrochemical detection using

Fig. 9 Sandwich enzyme immunoassay steps. Ab, antibody; Ag,

antigen; Ab*, enzyme-labeled secondary antibody; S, substrate; P,

product; and shaded oval, nonspecific binding blocker.


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the principles described above have been commercialized.92

Single cartridges for cardiac markers creatine kinase MB

(CK-MB), cardiac troponin I (cTnI) and B-type natriuretic

peptide (BNP) use alkaline phosphatase as the enzyme

label.

Affinity biosensors using impedance spectroscopy with gold

(Au) nanoparticles as the solid support for the biorecognition

element have been developed for the IgE antibody to a protein

allergen from dust mites,93,94 human immunoglobulin (hIgG),95

and carcinoembryonic antigen (CEA),96 a glycoprotein that is

produced only during fetal development. Au nanoparticles of

several different sizes are now commercially available and their

use in biosensors has become very popular.32 These Au

nanoparticles are also biocompatible. Biomolecules immobi-

lized on Au nanoparticles are usually stable and able to retain

their biological activity. Au nanoparticles are typically used to

form a single layer or a three-dimensional network on a

conductive electrode surface or are incorporated into a ceramic

sol–gel or polymer film.32

Impedance sensors using carbon nanotubes (CNTs) as the

sensor interface on which the capture Ab is immobilized have

also been reported.97–99 CNTs contain allotropes of carbon

arranged in sheets that have been rolled up into highly

conductive, hollow tubes of various nanometre dimensions.

CNTs have been incorporated in the sensing layer of

impedance biosensors due to their exceptionally high con-

ductivity and increased active surface area. CNT towers have

been used in impedance detection of mouse IgG and prostate

cancer cells.98,99

3.2 DNA hybridization biosensors

3.2.1 Background. Nucleic acid layers can also be used as

the biorecognition element coupled with electrochemical trans-

ducers in affinity biosensors. Electrochemical DNA hybridi-

zation biosensors are useful in the diagnosis of genetic or

infectious diseases, in environmental monitoring, to detect

microorganism contaminants in food and beverages, and for

national defense applications, among others.5

3.2.2 Detection mechanism. Complementary DNA base-

pairing is the basis for the biorecognition process in hybridi-

zation biosensors. Short, 20–40 basepair single-stranded DNA

segments with the ability to selectively bind with target analyte

are immobilized on the electrode surface.5 The DNA frag-

ments have to be immobilized in a way that retains their

stability, reactivity, accessibility to target analyte and optimal

orientation.5 Sensor surface coverage by DNA probes is also

important in minimizing nonspecific binding.5,100 An electrical

signal is produced when target DNA binds to the comple-

mentary sequence of the capture or probe DNA in a process

called hybridization. An electrochemical signal can result from

an electroactive indicator that binds preferentially to the DNA

duplexes instead of single-stranded DNA probes such as

ferrocenyl naphthalene diimide (FND).100 Electrochemical

measurement of a catalytic product from a captured enzyme

label such as horseradish peroxidase or alkaline phosphatase

can also be used as a measure of hybridization.101,102 The

enzymatic amplification of the binding event allows

measurements down to 3000 copies of target DNA or

zmols.103 Nanoparticle labels such as colloidal gold have also

been used to quantitate binding.2 Label-free electrochemical

measurement of hybridization induced changes in capacitance

or conductivity at the transducer surface have been used.5 The

nucleotide base guanine can be oxidized at the electrode and

the signal amplified by a redox mediator such as Ru(bpy)32+.

Like other biological macromolecules with complex

structures, the experimental conditions, such as tempera-

ture, ionic strength, and time allowed for hybridization, have

to be controlled in order to achieve high selectivity and

sensitivity.

3.2.3 Aptamer production. Single-stranded, 15–40 bases

long DNA or RNA oligonucleotide sequences that are used

as the biorecognition component called aptamers in biosensors

are rapidly screened in the SELEX (systematic evolution of

ligands by exponential enrichment) process for their ability to

selectively bind low molecular weight organic, inorganic or

protein targets.104,105 In solution, the synthetic nucleotide

chains form intramolecular interactions that fold the aptamer

molecules into a complex three-dimensional shape. The unique

shape of the aptamer allows it to bind tightly and selectively

with its target molecule. Aptamers can either bind to small

sections of macromolecules, such as proteins, or they can

engulf a small molecular target.

The selection process for aptamers has been around since

1990.104,105 An aptamer for a desired target molecule is chosen

from a large pool of random DNA and RNA sequences

generated using automated oligonucleotide synthesis methods

by successive cycles of binding to the immobilized target

molecule, followed by removing unbound material, and repli-

cating the bound nucleic acid strands for another round of

SELEX using the polymerase chain reaction (PCR). Chosen

aptamers after several cycles of SELEX can also be chemically

modified to increase their stability and affinity for a target

molecule. Once the sequence of nucleic acids in an aptamer for

a specific target is known, the aptamer can be synthesized in

large quantities. Like other biological molecules, aptamers

are sensitive to their environment and have to be protected

from high temperatures and DNAase enzymes. A variety of

strategies for developing aptamer-based electrochemical bio-

sensors are possible.106

3.2.4 Applications of DNA sensors. Osmetech has commer-

cialized an electrochemical sensor (eSensors

) based on the

selective reaction between a DNA capture probe immobilized

on the electrode surface and target DNA in the sample.107,108

The biosensor uses a sandwich type assay as shown in

Fig. 10A. Self-assembled monolayer (SAM) technology is used

to create the chemical layer attached to the gold electrode. The

monolayers are mixed SAMs, each comprised of a sequence-

specific capture probe (or probes) and an insulator com-

ponent. The DNA capture probe is immobilized on the gold

using an alkane thiol linker that projects it beyond a layer of

shorter alkane thiols. The shorter layer covers the surface

between the DNA capture probes and thereby minimizes

interference from redox active materials in the sample and

nonspecific adsorption, by blocking their access. Exposing the


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Fig. 10 Commercially available electrochemical DNA sensor (eSensors) by Osmetech: (a) detection principle, (b) assay genotyping principle, (c)

disposable biosensor printed circuit. (Published with permission of copyright holder, Clinical Micro Sensors, Inc. dba Osmetech Molecular

Diagnostics.)


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electrode to the sample results in hybridization between the

capture probe and the complementary strand of the target

DNA. The capture probe is designed to be shorter than the

complementary target strand, leaving a segment on the target

DNA where a signal probe containing an electroactive label

can bind. The label, ferrocene, is detected by measuring the

peak current for its oxidation by a positive potential scan in ac

voltammetry. The layer of alkane thiols is sufficiently thin as to

not interfere with the electrochemistry. The current is propor-

tional to the target DNA concentration in the sample. As

shown in Fig. 10B, genotyping can be done using different

ferrocene labels with distinguishable electrochemical poten-

tials for each label. The biochip consists of a microarray of

72 working electrodes, a Ag/AgCl reference electrode and two

auxiliary electrodes (Fig. 10C). Each working electrode of the

array can be interrogated independently which allows multiple

measurements to be made on the same chip. The chip is used

with a cartridge that features an auto-fill sample chamber,

microfluidic circulation to accelerate hybridization, and con-

tact with a resistive heating element. Osmetech has received

FDA clearance for eSensors

assays for detecting cystic fibrosis

carriers, and for identifying single-nucleotide polymorphisms

(SNPs) which result in increased sensitivity to warfarin, a

commonly prescribed blood anticoagulant.

3.3 Biosensors based on receptors

Receptors are proteins embedded in the cellular membrane

that specifically bind to their target analytes resulting in

physiological changes. The physiological response can be

opening ion-channels, producing second messenger systems,

or activating enzymes.1 A binding event at the receptor usually

causes the conformation of the receptor to change, which is

translated into an amplified electrochemical potential change.5

Unlike Abs that bind tightly with their complementary Ag,

receptors are like messengers that transmit signals upon ligand

binding between different parts of a biological system.1 Most

receptors are difficult to isolate and tend to bind to classes of

compounds having common chemical properties rather than

being highly specific for a given analyte like Abs.1,2 Therefore,

receptor-based biosensors are usually class-specific affinity

sensors that may not be a good feature for some biosensor

applications. Examples of receptor-based sensors include ion-

channel sensors where receptors in a lipid bilayer open or close

in response to a binding event with a ligand resulting in a rapid

ion flux through the membrane protein that causes a change in

the transmembrane conduction.109 The ion-channel membrane

proteins contain pores that allow ions such as Na+, K+ or

Ca2+ to flow through the channel until the potential difference

reaches equilibrium or the channel closes in response to a

stimulus. Also, nerve fibers from crayfish have been used to

monitor for local anesthetics and other drugs at low levels

(down to 10�15 M) with fast response times.110 Unfortunately,

these systems relying on axons from crayfish have a lifetime of

only 4 to 8 hours.

4. Conclusions

Catalytic and affinity biosensors with electrochemical detec-

tion continue to play an important role in many clinical,

environmental, industrial, pharmaceutical, defense, and

security applications due to their superior sensitivity and

selectivity. Although many electrochemical sensors are still

in the development and testing phases, some have reached

the consumer market as handheld devices, portable units used

for field measurements or are routinely used in a laboratory

setting. Recent developments in nanotechnology and mate-

rial science as well as being able to custom engineer the

biorecognition component will further push the develop-

ment of useful and reliable biosensor devices. The sometimes

limited shelf life and stability of the biorecognition component

as well as nonspecific binding continue to be the biggest

limitations of biosensors. However, many strategies have

helped with overcoming or minimizing these problems.

Notes and references

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