Electrochemical biosensors

CRITICAL REVIEW

www.rsc.org/csr | Chemical Society Reviews

Electrochemical biosensors Niina J. Ronkainen,*a H. Brian Halsallb and William R. Heinemanb 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 biological 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 biological recognition processes. These devices contain a biological

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

Department 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 b Department 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

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 Biocatalytic devices incorporate enzymes, whole cells or tissue slices that recognize the target analyte and produce electroactive 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,

Niina J. Ronkainen received her BS in chemistry and biology at Butler University (Indianapolis, USA) in 1997 and her PhD at the University of Cincinnati (USA) in 2003 where she specialized in bioanalytical chemistry. From 2003–2004 she taught chemistry as a visiting assistant professor at Tulane University (New Orleans, USA). In 2004 she joined Benedictine University as an assistant professor of chemistry. She currently Niina J. Ronkainen does basic research in biosensors and electrochemistry. She is an active member of the Chemical Education division of the American Chemical Society.

H. Brian Halsall is a professor of chemistry, and a member of the Sensors & Biosensors Group in the Department of Chemistry at the University of Cincinnati. He received a BSc (Hons) and PhD in chemistry at the University of Birmingham, UK. This was followed by postdoctoral work at UCLA, after which he joined the staff of the MAN Program at Oak Ridge National Laboratory before settling in Cincinnati. His principal research interests include biosensors, electrochemical immunoassay, and glycoprotein biochemistry.

a

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The Royal Society of Chemistry 2010

H. Brian Halsall

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nucleic acid, or a receptor. Immunosensors and DNA hybridization 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 biorecognition 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 application 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

William R. Heineman is a Distinguished Research Professor in the Department of Chemistry at the University of Cincinnati. He received a BS in chemistry at Texas Tech University and a PhD at the University of North Carolina in Chapel Hill and was a postdoctoral associate at Case Western Reserve University and The Ohio State University. His research interests include spectroelectrochemistry, electrochemical immunoassay, William R. Heineman sensors, and bioanalytical chemistry. He is a recipient of the Charles N. Reilley Award in Electroanalytical Chemistry and the Torbern Bergman Medal from the Analytical Section of the Swedish Chemical Society. 1748 | Chem. Soc. Rev., 2010, 39, 1747–1763

other detectable outcome.2 Enzymes, globular proteins composed 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. Parameters such as origin and availability of the biological component, 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 measurements 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 substances in a complex mixture, such as urine or blood, very selectively. This removes the need for time-consuming, laborintensive, 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 substrate 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 This journal is

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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 immunosensor could be developed to detect for that substance. Immunosensors are well known among analytical methods for their extremely low detection limits.6 Immunoassays and immunosensors have been developed for both quantitative and qualitative 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 environmental 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 biotechnology 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 biorecognition 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 transducer 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 potential (potentiometry) or alters the conductive properties of the medium between electrodes (conductometry).3 Use of electrochemical impedance spectroscopy by monitoring both resistance 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 This journal is

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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 detection. 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 amperometric 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 amperometry 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 Chem. Soc. Rev., 2010, 39, 1747–1763 | 1749

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 biosensors. 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 significant advantage when the sample sizes are limited.21,22 Furthermore, 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 detection 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 1750 | Chem. Soc. Rev., 2010, 39, 1747–1763

Fig. 2 Diagram of a screen printed electrode (SPE). Ref., reference electrode; Aux., auxiliary electrode; and Work., working electrode.

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. 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 perturbation 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, respectively. 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 electrochemical 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 labelfree 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. However, regenerating the sensing surface for a subsequent measurement 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

Fig. 4 A diagram of an Ab–Ag affinity sensor with impedimetric detection.

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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 carefully chosen. Being directly able to impedimetrically monitor the Ab–Ag binding helps by-pass the aforementioned limitations. EIS is also insensitive to most environmental disturbances. 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 electrochemical 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 spectroscopy 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 reaction. 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 immunosensor that is based on an electrochemical sandwich immunoassay, 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 electrode 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. Chem. Soc. Rev., 2010, 39, 1747–1763 | 1751

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. Miniaturization 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 microfluidic 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 microelectrode 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 catecholamine 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 electrochemical 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

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 enzymebased 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 periodically since it gradually loses activity. Also, clever electrochemical 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 applications, 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 implemented by Updike and Hicks, who constructed the first functional biocatalytic sensor for glucose.51 In the first-generation biosensors, an oxidase enzyme is immobilized behind a semipermeable 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)

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. Alternatively, 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 1752 | Chem. Soc. Rev., 2010, 39, 1747–1763

Fig. 5 Oxygen-dependent first-generation biosensor with amperometric 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+ 2MediatorRed - 2MediatorOx + 2e

(4) (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 This journal is

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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 dehydrogenases to be used as the biorecognition element in the sensor thereby expanding the list of possible target analytes. Third-generation biosensors have the biorecognition component 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 measurements since neither mediator nor enzyme need to be added. This self-contained nature also lowers the cost per measurement 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 transducer’s sensing surface without blocking the active site of the enzyme or drastically altering the enzyme geometry.2 Immobilization methods are considered successful if the biosensors prepared 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 proportional to substrate concentration for reliable quantitation of the analyte. Km, the Michaelis constant, is the substrate concentration 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 immobilized 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 selectivity 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 adsorption are primarily very weak van der Waals forces with occasional hydrogen bonds that are not very stable or permanent.1

Fig. 7 Common methods of immobilizing enzymes onto an electrode surface.

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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 immobilized 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 substrate 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 components 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 applications 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 technology. The test strip also contains two electrodes, one holding the enzyme and a mediator for the amperometric detection This journal is

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of glucose, and the other serving as a reference electrode. The current produced when a potential is applied gives a readout to a liquid crystal display on the glucose meter. The commercially sold glucose test strips are second- or thirdgeneration 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

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 indicators 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

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 electrophoresis 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 intermittent ‘‘spot’’ lactate monitoring are being manufactured by This journal is

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Interference-based enzyme electrodes are probes used for quantitative 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 biocatalytic 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 Chem. Soc. Rev., 2010, 39, 1747–1763 | 1755

Live microorganisms have also been coupled with electrochemical transducers (i.e. electrodes) to monitor biotechnological processes such as brewing, food manufacturing, waste-water treatment, energy production, and pharmaceutical 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 components, 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 instrumentation. 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 common, the development of more convenient immunosensors that are easier and faster to use gained momentum. Most applications of immunoassays (IA) and immunosensors 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 radioimmunoassays were developed for biomedical applications such as detecting hormones and disease related proteins, but applications 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 chromatography 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 effective. The immunoassays and affinity biosensors are relatively 1756 | Chem. Soc. Rev., 2010, 39, 1747–1763

easy to use once fully developed and optimized, and the reductions in chemicals used, waste disposal, expensive instruments 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 E 150 kDa, are produced by a host in response to the presence of a foreign molecule called antigen (Ag).5 Antigens are anything 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 recognized 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 lightchain 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 noncovalent 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 noncovalent 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 distances. Sometimes the exceptional selectivity can be a disadvantage 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 This journal is

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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 crossreactivity 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 interference for the affinity sensor.

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

3.1.4 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 misoriented 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 immobilization methods include biotin–streptavidin linkages,6,22,29 adsorption to a conductive polymer matrix such as polypyrrole,81 and covalent binding.3,82

3.1.2 Antibody production. The production of the antibodies (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. Sometimes 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 monoclonal 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. Polyclonal 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

3.1.5 Formats for enzyme immunoassays. Enzyme immunoassays (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. However, 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 quantitation 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 immunosorbent 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 immunoreactants 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. 1758 | Chem. Soc. Rev., 2010, 39, 1747–1763

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.

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 immunoassays 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. Selfassembling 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 reagents 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 entropically 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 Detergents 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 instruments 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 containing a microfabricated biosensor array. The system automatically 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 This journal is

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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 immobilized 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 conductivity 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 transducers in affinity biosensors. Electrochemical DNA hybridization 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 basepairing is the basis for the biorecognition process in hybridization 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 fragments 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 complementary 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 This journal is

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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 temperature, 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 replicating 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 biosensors are possible.106 3.2.4 Applications of DNA sensors. Osmetech has commers cialized an electrochemical sensor (eSensor ) 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 sequencespecific capture probe (or probes) and an insulator component. 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 Chem. Soc. Rev., 2010, 39, 1747–1763 | 1759

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 proportional to the target DNA concentration in the sample. As shown in Fig. 10B, genotyping can be done using different ferrocene labels with distinguishable electrochemical potentials 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 contact with a resistive heating element. Osmetech has received s FDA clearance for eSensor 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 ionchannel 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 detection continue to play an important role in many clinical, This journal is

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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 material science as well as being able to custom engineer the biorecognition component will further push the development 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.

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