Prospects of conducting polymers in biosensors

Analytica Chimica Acta 578 (2006) 59–74 Prospects of conducting polymers in biosensors Bansi D. Malhotra a,∗ , Asha Chaubey b , S.P. Singh a a Biomolecular Electronics and Conducting Polymer Research Group, National Physical Laboratory, Dr. K.S. Krishnan Marg, New Delhi 110012, India b Regional Research Laboratory, Jammu, India Received 18 February 2006; received in revised form 17 April 2006; accepted 20 April 2006 Available online 29 April 2006 Abstract Applications of conducting polymers to biosensors have recently aroused much interest. This is because these molecular electronic materials offer control of different parameters such as polymer layer thickness, electrical properties and bio-reagent loading, etc. Moreover, conducting polymer based biosensors are likely to cater to the pressing requirements such as biocompatibility, possibility of in vivo sensing, continuous monitoring of drugs or metabolites, multi-parametric assays, miniaturization and high information density. This paper deals with the emerging trends in conducting polymer based biosensors during the last about 5 years. © 2006 Elsevier B.V. All rights reserved. Keywords: Conducting polymer; Biosesnor; Soliton; Polaron; Bipolaron; Enzyme; Glucose oxidase; Urease; Cholesterol oxidase; Deoxyribonucleic acid (DNA); Peptide nucleic acid (PNA); Microorganism; Amperometric; Potentiometric; Immunosensor 1. Introduction Biosensors have recently attracted much interest. This is because these interesting bio-devices have been shown to have applications in clinical diagnostics, environmental monitoring, food freshness and bioprocess monitoring [1–249]. A number of materials such as polymers, sol–gels and conducting polymers have been used to improve the stability of the biomolecules used in the fabrication of the desired biosensors. In this context, polymers have become the materials of choice for recent technological advances in biotechnology. Initially polymers were thought to be insulators and were utilized in various electrical and electronic devices. The discovery of poly (sulphur nitride) [(SN)x ] which becomes superconducting at low temperatures led to the renewed interest in polymers [10]. The electrical conductivity of [(SN)x ] can be enhanced by several orders, i.e. 105 S cm−1 by simple doping with oxidizing agents, e.g. I2 , AsF5 , NOPF6 (p-doping) or reducing agents (n-doping), e.g. sodium napthalide. Subsequently, a new class of polymers such as poly-para-phenylene (PPP), polyphenylene sulphide (PPS), polythiophenes and polypyrroles (PPy) were reported ∗ Corresponding author. Tel.: +91 11 25734273; fax: +91 11 25726938. E-mail address: [email protected] (B.D. Malhotra). 0003-2670/$ – see front matter © 2006 Elsevier B.V. All rights reserved. doi:10.1016/j.aca.2006.04.055 [11–13]. Diaz et al. produced coherent films of PPy with conductivity of 100 S cm−1 and this conducting polymer exhibits excellent air stability [14]. Chen et al. have estimated the concentration of cytochrome C using electrochemically prepared conducting polymers based on ferrocene substituted thiophene and terthiophene [44]. Many other conducting polymers such as poly(3,4-ethylenedioxythiophene) (PEDOT), polyfuran, polyindole, polycarbazole, polyaniline, etc. have been synthesized and studied extensively [12–16,45]. Compared to saturated polymers, conducting polymers have different electronic structures. Chemical bonding in conducting polymers provides one unpaired electron, i.e. ␲ electron per carbon atom in the backbone of the polymer. Carbon atoms are in sp2 pz configuration in ␲ bonding and orbitals of successive carbon atoms overlap providing delocalization of electrons along the backbone of polymer [12]. This delocalization provides the charge mobility along the backbone of the polymer chain and induces unusual properties such as electrical conductivity, low ionization potential, low energy optical transitions and high electron affinity. The ␲ bonds in conjugated polymers are highly susceptible to chemical or electrochemical oxidation or reduction. The origin of electrical conduction in conducting polymers has been ascribed to the formation of non-linear defects such as solitons, polarons or bipolarons formed either during doping or polymerization of a monomer [19,46,47,57]. 60 B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 The conductive and semiconducting properties of these polymers make them an important class of materials for a wide range of electronic, optoelectronic and biotechnological applications such as in rechargeable batteries, molecular electronics, electronic displays, solar cells, ion exchange membrane in fuel cells, diodes, capacitors, field-effect-transistors, printed circuit boards, chemical sensors, drug release systems and biosensors, etc. [14–17]. It is being projected that conducting polymers can be used to transport small electronic signals in the body, i.e. act as artificial nerves. Perhaps modifications to the brain may eventually be contemplated. Scientists have used films in a neurotransmitter as a drug release system into the brain [18]. Conducting polymers have emerged as potential candidates for biosensors. Gerard et al. have reviewed the literature on applications of conducting polymers to biosensors [57]. Geetha et al. [77] have discussed the applications of conducting polypyrrole to drug delivery. Andreescu and Sadik have reviewed the challenges and trends in biosensors for environmental and clinical monitoring [78]. Cosnier [79,80] has discussed the analytical applications of affinity biosensors based on electropolymerized films. Ramaniviciene and Ramanavicius have reported an interesting overview on the potential use of conducting polymers as electrochemical based affinity biosensors [81]. Malinauskas et al. have reviewed the electrochemical aspects of conducting polymer-based nano-structured materials for application to super-capacitors, energy conversion systems, batteries and sensors [82]. Wanekaya et al. have reviewed recent advances in biosensors based on one-dimensional (1-D) nanostructures [83]. Ming-Hung Lee et al. have highlighted the current developments of DNA-based bioanalytical microsystesm for point-of-care diagnostics [84]. Adhikari and Majumdar have discussed the role of non-conducting and intrinsically conducting polymers in sensor devices [56]. Terry et al. [86] have assessed the future and current trends of biosensors in food industry. Drummond et al. have discussed numerous approaches to electrochemical detection based on modified electrodes, electrochemical amplifications with nanoparticles and electrochemical devices using DNA-mediated charge transport chemistry and electrochemistry of DNA-specific redox reporters [87]. Habermuller et al. have reported on the various electron-transfer mechanisms operating in amperometric biosensors [88]. Kerman et al. have predicted that electrochemical DNA biosensors with suitable microfabrication techniques are likely to be increasingly popular in the near future [89]. The present paper focuses on the prospective applications of conducting polymers in biosensors. 2. Importance of conducting polymers to biosensors Conducting polymers [48–114] have attracted much interest as suitable matrices of biomolecules and can be used to enhance stability, speed, sensitivity and hence are finding increasing use in medical diagnostics [20–23,51–96]. A number of techniques such as physical adsorption, electrochemical entrapment and covalent attachment based on ethyl-dimethylaminopropylcarbodiimide (EDC) and N-hydroxy-succinimide (NHS) coupling chemistry, have been used to improve the sta- bility of the desired biomolecules onto conducting polymers [57,80,170,171,242–245]. Schuhmann et al. reported that functionalization of conducting polypyrrole films provides suitable surface for covalent linkage of enzymes after carbodiimide activation [224]. Cosnier et al. have discovered that oxidative electropolymerization of a new (dicarbazole) derivative functionalized by N-hydroxysuccinimide group in acetonitrile results in electroactive poly (dicarbazole) films. Chemical functionalization of the poly (dicarbazole) film could be easily performed by successive immersions in aqueous and mediator solutions [170]. Devouge et al. have reported a practical molecular clip for immobilization of receptors and biomolecules on device surface based on the photographting of o-succinimidyl-4-(pazido-phenyl)butanoate [241]. Liu et al. have constructed a bio-electrode in which a self-assembled monolayer containing a novel norbornylogous bridge by covalently attaching to flavin adenine dinucleotide (FAD), the redox active centre of several oxidase enzymes [243]. It has been proposed that a decrease in the electronic coupling between the redox active FAD and the electrode following reconstitution of the glucose oxidase is responsible for the inability of the enzyme to be turned over under anaerobic conditions. The techniques of incorporating biomolecules into electrodepositable conducting polymeric films permit the localization of biologically active molecules on electrodes of any size or geometry and are particularly appropriate for the fabrication of multi-analyte micro-amperometric biosensors [24]. Conducting polymers can act as transducers in biosensors. A transducer converts a biochemical signal resulting through the interaction of a biological component into an electronic signal. Suitable transducing systems can be adapted in a sensor assembly depending on the nature of the biochemical interaction with the species of interest. The physical transducers vary from electrochemical, spectroscopic, thermal, piezoelectric and surface wave technology [25,26]. The most common electrochemical transducers being utilized are amperometric and potentiometric. An amperometric biosensor measures the current produced during the oxidation or reduction of a product or reactant at a constant applied potential. Such sensors have fast response times and good sensitivities. However, the excellent specificity of a biological component can be compromised by the partial selectivity of the bio-electrode. This lack of specificity requires sample preparation, separation or some compensation for interfacing signals. Potentiometric biosensors relate electrical potential to the concentration of analytes by using ion-selective electrodes or gas-sensing electrodes as the physical transducers [27,28]. Electrically conducting polymers are known to have considerable flexibility in chemical structures, which can be modified as required. By chemical modeling and synthesis, it is possible to modulate the required electronic and mechanical properties of conducting polymers. Moreover, the polymer itself can be modified to bind protein molecules [29–31,116]. Another advantage offered by conducting polymers is that the electrochemical synthesis allows direct deposition of a polymer on the electrode surface while simultaneously trapping the protein molecules [32,33]. It is thus possible to control the spatial distribution of B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 the immobilized enzymes, the film thickness and modulation of enzyme activity by changing the state of the polymer. The development of a technology in this field heavily depends on the understanding of interactions at the molecular level between the biologically active protein, either as a simple composite or through chemical grafting. For the proper relay of electrons from the surface of an electrode to the enzyme active site, the concept of ‘electric wiring’ has been reported [34,35]. Conductive polymers can be reversibly doped and undoped using electrochemical techniques accompanied by significant changes in conductivity. Besides this, the optical properties of these films can be used as a signal for investigating biochemical reactions [36,37]. The electrical conductivity of conducting polymers changes over several orders of magnitude in response to change in pH, redox potential or their environment [38]. Conducting polymers have the ability to efficiently transfer electric charge produced by a biochemical reaction [39]. Moreover, conducting polymers can be deposited over defined areas of electrode. This unique property of conducting polymers along with the possibility to entrap enzymes during electrochemical polymerization has been exploited for the fabrication of amperometric biosensors [40–52]. Besides this, conducting polymers exhibit exchange and size exclusion properties due to which these are highly sensitive and specific towards desired substrates [53–55]. Bioelectrochemistry and electroanalysis of biologically important substances are being intensively studied for application as chemically modified electrodes [57–60]. Conducting polymers are also known for their ability to be compatible with biological molecules in neutral aqueous solutions [59]. Wrighton et al. demonstrated sensor response to oxidizing and reducing species in solution [60]. These sensors were based on changes in the electrical conductivity that accompanies oxidation and reduction of polymers such as polypyrrole, poly-(3-methylthiophene), and polyaniline [61–63]. The electrochemical conversion is known to occur at a conducting polymer/solution interface when the overall rate of both the biochemical reaction and charge transport exceeds substantially the rate of mass transport of solution species to the electrode. This phenomenon is a diffusion controlled electrocatalytic process. The combination of relative rates for the three processes favors an efficient, reversible anodic oxidation of analyte and can be used for the amperometric detection at relatively low electrode potential. Redox mediators can be effective in the technical development of biosensors. The redox-active sites in redox proteins are generally shielded by insulating peptides and sugars. The electron transfer between enzymes and conducting polymers (or electrodes) can be accelerated by using small electron transfer mediators or shuttles, such as quinones and ferrocenes, which enter the clefts of proteins and shuttle between the redox sites of enzymes and electrode surfaces [64]. Gorton et al. have discussed the role of direct electron transfer between hemecontaining enzymes and electrodes as the basis of third generation biosensors [65]. Schuhmann has described biosensors based on (i) electron-transfer via conducting polymers, (ii) electrontransfer cascades via redox hydrogels, (iii) anisotropic orienta- 61 tion of redox proteins and (iv) direct electron transfer between redox proteins and electrodes modified with self-assembled monolayers [66]. The interest in nanomaterials for biosensing applications has recently emerged [233–239]. Nanostructures are inorganic, organic or composite materials synthesized in various forms with sizes down to the nanometer. The size dependent changes in the physical and chemical properties of materials make them different than that of their bulk counterpart. The change in physical properties such as electrical, electronic, magnetic, optic and thermodynamic properties makes nanomaterials important candidate for technological development of biodevices. The large surface to volume ratio provides substantial changes in chemical properties. When size of these materials is of the order of de Broglie wavelength, quantum confinement effect becomes prominent which also gives rise to the unusual electrical and optical properties. The examples of materials falling in this category (with different molecular symmetries) are fullerenes and carbon nanotubes. Like conjugated polymers, molecular structure of fullerenes and carbon nanotubes also consist of number of C C double and single bond. The fullerenes and carbon nanotubes can be termed as a class of conjugated polymers. Semiconducting nature and possibility of p-type and n-type doping which can be made possible due to the presence of conjugated carbon structure with unusual molecular symmetries makes them compatible with conducting polymers and promising materials from technological point of view. Biological molecules can be covalently attached to the carbon nanotubes. Single wall and multi-wall carbon nanotubes were made water soluble via esterification of nanotube bound carboxylic acid by oligomeric polyethylene glycols which were then biofunctionalized by bovine serum albumin [68,69]. The functionalization of carbon nanotubes with bovine serum albumin (BSA) has been found to be highly water soluble. The majorities of carbon nanotubes are intimately associated with the BSA protein and remain bioactive. This method may be used to introduce carbon nanotubes into other biological and biomedically important systems [68]. Ramanathan et al. have functionalized single wall carbon nanotubes with amino groups, which could further be covalently attached to polymers or biological systems such as DNA and carbohydrates [67]. Varfolomeyev et al. have reviewed the recent developments in electrochemistry and electro-analytical chemistry of carbon nanotubes as new nanomaterials [68]. Unique structural property of CNTS makes these materials attractive to accommodate electrochemically or biologically functional elements, e.g. quantum dots, organic compounds, biomaterials like enzymes, DNA, proteins and antibodies paving the way for the fabrication of highly sensitive and stable biosensors. Ramanathan et al. have fabricated bioaffinity sensors using biologically functionalized conducting polymer nanowire [225,226]. It has been shown that sensitivity of potentially single-molecule detection can be achieved by adjusting the nanowire’s conductivity closer to the lower end of a semiconductor. It has been demonstrated that similar to biological modification of PPy nanowires, monomers such as thiophene and aniline that can be electropolymerized from an aqueous environment can also be used. Carrara et al. have uti- 62 B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 lized improved nanocomposite materials of poly(o-anisidine) (POA) containing titanium dioxide nanoparticles (TiO2 ), carbon black and multi-walled carbon nanotubes (MWNT) for biosensing applications based on electrochemical impedance spectroscopy [232]. The biotin-labeled biomolecules can be immobilized on the avidin-modified electrode surface through avidin–biotin complexation [104,136,216,246–248]. Nzai et al. have described various techniques for the surface derivatization within biotin and avidin and for the coupling of the enzymes [246]. It is proposed that the possibility of constructing protein architecture is based on the non-covalent interaction of avidin and biotin. Nobs et al. have covalently bound NeutrAvidinTM to the surface of poly (dl-lactic acid) (PLA) nanoparticles with the aim of attaching targeting compounds such as proteins to their surface [247]. Their studies indicate that other proteins such as antibodies could be coupled to the nanoparticles for active targeting. Furthermore, PLA nanoparticles are interesting candidates for active targeting with biotinylated antibodies using the biotin–avidin interaction in a two-step procedure. Gref et al. have recently shown that biotin-poly (ethylene glycol)-poly (␧-caprolactone) (B-PEG-PCL) can be helpful for studying the interaction between cells and functionalized nanoparticles with surface characteristics (ligand type and density, PEG layer density and thickness) [248]. The impedance spectroscopy has been used to fabricate a number of conducting polymer and nanoparticles biosensors based on enzymes, antibodies and mico-organisms, respectively. Li et al. have recently reviewed the literature on impedimetric biosensors [69]. Heeger and Heeger have revealed that conducting polymer luminescence can be quantitatively manipulated to fabricate a variety of real-time biosensing applications including medical diagnostics and toxicology [70]. Molecular imprinting is gradually becoming a versatile technique for the preparation of artificial receptors based on molecularly imprinted conducting polymers (MIPs) containing tailormade sites [74]. Molecular imprinting technology can be used for the manufacture of synthetic polymers with pre-determined molecular recognition properties. MIPS are highly stable synthetic polymers having molecular recognition properties due to cavities in the polymer matrix that are complementary to the analyte (ligand) both in shape and the positioning of the functional groups and hence have been used to obtain stabilized biological response [75]. Some of these polymers exhibit very high affinity constants and selectivity comparable to naturally occurring molecular recognition systems such as antibodies. Over-oxidized polypyrrole shows improved selectivity due to oxygen functionality signed to the removal of positive ions from PPy films. Sensors based on methyl pyrrole (mPPy) have been fabricated for detection of 1-naphthalenesulphonate and 1-serotonin [75,76]. Ho et al. have utilized poly(3,4ethylenedioxythiophene) (PEDOT) to immobilize MIP particles of morphine onto the indium–tin-oxide (ITO) glass. The sensitivity, detection limit, linear range and the signal/noise ratio of the modified morphine MIP/PEDOT electrode were experimentally estimated as 41.63 ␮A cm−2 mM−1 , 0.3 mM, 0.1–2 mM and 3, respectively [76]. Fig. 1. Schematic of a conducting polymer based biosensor. Fig. 1 describes the schematic of a conducting polymer based biosensor. 3. Conducting polymer based biosensors 3.1. Conducting polymer based enzyme biosensors Enzymatic biosensors utilize the biospecificity of an enzymatic reaction, along with an electrode reaction that generates an electric current or potential difference for quantitative analysis [89–175]. The biomolecules such as glucose, cholesterol, urea, triglycerides, creatinines, pesticides are important analytes due to their adverse effects on health. Enzymatic biosensors utilize the biochemical reactions, i.e. analyte and enzyme resulting in a product (hydrogen/hydrogen peroxide/hydroxyl/ammonium ion) that can be detected and quantified using a transducer (amperometric/potentiometric/optical thermal/piezoelectric). In general, many oxidoreductases including glucose oxidase catalyze the oxidation of substrates by electron transfer to oxygen to form hydrogen peroxide. These oxidoreductase enzymes can be immobilized on conducting polymer surfaces and the H2 O2 formed as a result of enzyme and the corresponding analytes may be measured amperometrically [98–110]. However, it has not been possible to discriminate between the direct electron transfer from the oxidation of hydrogen peroxide at polymer surface and that at the underlying electrode [99–111]. Since conducting polymers are insoluble in aqueous solutions, electropolymerization has been frequently used to create a matrix for immobilization of enzymes at the electrode surface, and the sensor response was obtained by the oxidation of hydrogen peroxide [112,113]. Belanger et al. suggested that the reaction of hydrogen peroxide and polypyrrole decreases the electrical conductivity [117]. Over potential can be reduced by using mediators; small molecules that shuttle between electrode and analyte to accelerate electron flows and the formal potential of the mediator should be close to or positive to that of the analyte. Tian et al. [131] have described an amperometric biosensor for the detection of H2 O2 based on horseradish peroxidase/polypyrrole (PPy) membrane deposited onto the surface of ferrocenecraboxylic acid mediated derived sol–gel derived composite carbon electrode. This biosensor had a linear range from 2 × 10−5 to 2.6 × 10−3 M), detection limit as 0.2 mM and retained its 90% of the initial sensitivity (52.2 ␮A mM−1 ) for about 1 week. It may be remarked that there are some dehydrogenases that may directly transfer B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 electrons to conducting polymers without producing hydrogen peroxide [223]. Sharma et al. have demonstrated that electrochemically prepared poly(aniline-co-fluoroaniline) films for immobilization of glucose oxidase using physical adsorption technique [94]. The shelf-life of these conducting polymer glucose electrodes was found to be about 15 days. Singhal et al. prepared poly(3-hexylthiophene) Langmuir–Blodgett films by simultaneous entrapment of glucose oxidase and transferred onto indium–tin-oxide glass [120]. This glucose electrode was found to have detection limit of 50 mg dL−1 and sensitivity of 0.75 nA mg−1 dL−1 and linearity from 100 to 500 mg dL−1 of glucose and stability of 75 days at 4 ◦ C [120]. Subsequently, these researchers reported a glucose biosensor based on Langmuir–Blodgett films of polyvinylcarbazole [121]. These glucose electrodes could be repeatedly used for 15 days for the estimation of glucose from 100 to 500 mg dL−1 and were stable for about 5 months at 4 ◦ C. Uang and Chou have investigated the pH effect on the electropolymerization in the presence of glucose oxidase on the characteristics of glucose oxidase/polypyrrole (PPy) biosensor. This glucose sensor had linearity from 0 to 10 mM, sensitivity as 7 nA mM−1 and could be used for about 2 weeks [123]. Borale et al. have constructed glucose biosensors based on electropolymerized polyaniline, poly (o-toluidine) and poly (anilinie-cotoluidine), respectively [124]. Sekar and Becerik have electrochemically entrapped glucose oxidase into polypyrrole matrix alongwith p-benzoquinone in piperazine ethyl sulphate (PIPES) buffer [125]. It is revealed that application of the artificial network analysis applied to the non-linear calibration plot can be used to predict the sensor failure detection. Ramanavicius et al. have prepared polypyrrole in the presence of glucose oxidase from Pencillum vitale, glucose and oxygen revealing that optimal conditions of glucose oxidase activity (pH 6.0) are similar to pyrrole polymerization reaction (pH 6.5) indicating possible biomedical applications of this conducting polymer [126]. These authors later reported the polypyrrole coated glucose oxidase nano-particles by self encapsulation of GOX in polypyrrole during polymerization for application as amperometric glucose biosensor and showed that the Km value of polypyrrole coated GOX is 10 times larger than that for the native GOX [166]. The shape/size and other properties of these polypyrrole coated glucose oxidase nano-particles have not yet been investigated. Pan et al. have reported CNT based glucose biosensor using immobilization of glucose oxidase in poly (o-aminophenol) and carbon nano-tube composite through electrochemical polymerization onto gold electrodes [118]. They have shown that this biosensor has detection limit of 0.01 M, sensitivity of 0.114 A cm−2 and is stable for about 30 days. These authors, however, did not investigate the thermal stability of this CNT glucose electrode. Curulli et al. have studied the influence of electrolyte nature and its concentration on the kinetics of electropolymerization of monomers such as 1,2- and 1,4-diaminobenzene, 2,3- and 1,8-diaminonaphthalene, o-anisidine and on the resulting morphology of nanotubules [133]. Glucose oxidase (GOx), lactate oxidase (LOD), l-aminoacid oxidase (l-AAOD), alcohol oxidase(AOD), glycerol-3-phosphate oxidase (GPO), lysine 63 oxidase (LyOx) and choline oxidase (ChOx) were immobilized on Prussian Blue (PB) layer supported on 1,2-diaminobenzene (1,2-DAB) nanotubes onto platinum electrodes. The nanostructured poly (1,2-DAB) based glucose biosensor showed detection limit of 5 × 10−5 mol L−1 , a wide linear range (5 × 10−5 to 5 × 10−3 mol L−1 ) and stability for 4 weeks at room temperature. Similar results were obtained for choline, l-leucine, l-(+)lactic acid, ethanol, lysine and glycerol-3-phoaphate, respectively. Callegavi et al. have fabricated the amperometric biosensors for lactate, phenol, catechol and ethanol by incorporation of their respective enzyme in carbon nanotubes (CNT) matrix [122]. They have shown that the use of CNT electrode yields better results as compared to normal carbon paste electrode. The observed stability of the H2 O2 sensor has been attributed to the nano-tubular structure. It was shown that these nano-structured polymers can be utilized for fabrication of enzyme biosensors. Trojanowicz and Miernik have employed avidin–biotin interactions for the immobilization of glucose oxidase on bilayer lipid membrane (BLM) formed on polypyrrole and polyo-phenylylenediamine) electrodeposited onto platinum wire, respectively [136]. A lipid solution comprising of 1.6 mg of biotin DHPE (N-(biotinoyl)-1,2-dihexadecaneoyal-sn-glycero3-phosphor ethanolo-amine, trethylammonium salt) and 0.4 mg of cholesterol dissolved in 1 cm3 of a mixture of n-decane and n-butanol (8:1, v/v) was used for lipid formation. The solid supported BLMs were formed on a bare or polymer-modified platinum surface (0.6 mm in diameter in each case) by immersing the tip of the electrode in the lipid solution for 5 min and then in avidin–enzyme conjugate for 5 min. These BLM based glucose biosensors resulted in stable and sensitive response with significant reduction from electro-active species such as ascorbic acid, cholesterol and uric acid, respectively. Mylar et al. have reported improved signal linearity of enzyme (glucose oxidase) ultra micro-electrodes fabricated via sonication and deposition of polysiloxane coating onto the working glucose oxidase/polyaniline electrode coated with insulating diaminobenzene [137]. He et al. [139] have developed a sensitive hydrogen peroxide probe based on amplified fluorescence quenching ability of poly(9,9-bis(6′ -N,N,Ntrimethylammonium-hexyl) fluorene phenylene (PFP-NMe3 + , a cation conjugated polymer) and peroxyfluor-1 borate protective groups. The hydrogen peroxide probe has a detection range of 15–600 nM and can be used for glucose detection. Conducting polymer based self-regulating insulin delivery system has recently been reported [140,141]. It has been brought out that conducting polypyrroles provide good alternatives due to biocompatibility and in vivo stability. Experiments are however being conducted to choose a suitable conducting polymers or composites for in vivo applications. Arslan et al. have fabricated a polysiloxane/polypyrrole/ tyrosinase electrode by entrapment of tyrosinase in conducting matrix by electrochemical copolymerization for determining phenolic content of green and black tea [128]. Compared to sodium dodecyl sulfate (SDS) doped electrodes, the p-toluene sulfonic acid (PTSA) doped enzyme electrodes show good pH resistance and temperature stability and the effect of supporting electrolyte (PTSA). However, the PTSA doped conducting 64 B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 polymers could not be used to detect phenolic components in tea because of the apparent high Km value. Vedrine et al. have fabricated an amperomteric tyrosinase biosensor based on conducting poly(3,4-ethylenedioxythiophene) (PEDOT) for estimation of herbicides and phenolic compounds [140]. The detection limits for monophenol and di-phenol ranged from 5 to 500 nM and the detection limit for diuron and attrazine were found to be 0.5 and 1 mg L−1 , respectively. Böyükbayram et al. have made a comparative study of immobilization methods of tyrosinase (Tyr) on electrolpoymerized conducting and non-conducting polymers for application to the detection of dichlorvos organaophosphorous insecticide [141]. A detection limit of 0.06 ␮m was obtained for dichlorvos using mediator 1,2-naphthoquinone-4-sulphonate (NQS) and Tyr. Jiang et al. have demonstrated that incorporation of polyvinylalcohol (PVA) onto a polymer film caused higher sensitivity than that of pure PPy sensor [130]. Rahman et al. have fabricated the amperometric biosensor for choline by immobilizing choline oxidase (ChOx) with and without horseradish peroxidase (HRP) onto carboxylated thiophene modified electrodes [135]. They have shown that the electrode with HRP has better performance and these biosensors have the response time of 5 s and selectivity to interfering species. Kan et al. have reported stronger affinity between uricase and polyaniline prepared by template process resulting in the increased stability of this polyaniline–uricase biosensor [134]. However, the activation energy of this uricase electrode has been found to be larger by 29.9 kJ mol−1 than the uricase–polyaniline biosensor fabricated using two-step processes. Haccoun et al. have fabricated a reagentless lactate biosensor using electro-copolymerized copolymer film of poly(5hydroxy-1,4-naphthoquinone-co-5-hydroxy-3-acetic acid-1,4naphthoquinone) [142]. These studies show that the presence of interferents like acetaminophen, glycine and ascorbic acid does not influence the response of this mediated (quinone group) copolymer electrode. However, this biosensor can be used for l-lactate estimation upto 1 mM of l-lactate. Bartlett et al. have found that poly(aniline)–poly(anion) composites films can be utilized for electrochemical oxidation of NADH at around 50 mV versus SCE and pH 7 [143]. These results have implications towards the technical development of microelectrodes, biofuel cells and amperometric biosensors. Chaubey et al. coimmobilized lactate oxidase (LOD) and lactate dehydrogenase (LDH) on electrochemically prepared polyaniline (PANI) films by physical adsorption technique [145]. The LOD/LDH/PANI electrodes were shown to have linearity from 0.1 to 1 mM of lactate and detection limit of 5 × 10−5 M and stability of 3 weeks at 4–10 ◦ C. Subsequently, these researchers electrochemically entrapped polyaniline onto sol–gel derived tetraethylorthosilicate (TEOS) films deposited onto indium–tin-oxide coated glass electrodes for immobilization of lactate dehydrogenase (LDH) [146]. The application of external layer of polyvinyl chloride (PVC) onto sol–gel/PANI/LDH electrodes resulted in the extended linearity from 1 to 10 mM of l-lactate. These electrodes had a shelf-life of 8 weeks at 0–4 ◦ C and the common interferents such as glucose (100 mg dL−1 ), uric acid (35 mg dL−1 ), glutamic acid (25 mg dL−1 ) and ascorbic acid (25 mg dL−1 ) affected the observed amperomteric response. Singh et al. have reported conducting polypyrrole/cholesterol oxidase electrodes based on conducting polypyrrole films for estimation of cholesterol ester concentrations of 1–8 mM [147,148]. The applicability of these electrodes to serum samples has not yet been demonstrated. These researchers have co-immobilized cholesterol oxidase, cholesterol esterase and peroxidase onto electrochemically prepared polyaniline films. This polyanline based cholesterol biosensor has a response time of 240 s, an apparent Km value as 75 mg dL−1 , sensitivity value as 0.042 ␮A mg dL−1 , detection limit as 25 mg dL−1 , shelf-life of 6 weeks and can be used to estimate cholesterol concentration up to 500 mg dL−1 . Asberg and Inganas have cross-linked horseradish peroxidase in highly conducting poly (3,4-ethylene dioxy thiophene) (PEDOT)/(polystyrene sulphonate) (PSS) dispersion using poly-4-vinylpyridine for estimation of hydrogen peroxide in the concentration range of 0–30 ␮M [149]. Grennam et al. have shown that horseradish peroxidase immobilized screen-printed electrodes based on chemically polymerized polyaniline/polyvinylsulphonate films can be used for the mass production of biosensors [150]. Morrin et al. have electrochemically applied nanoparticulate polyaniline (PANI) doped with dodecylbenzenesulphonic acid (DBSA) to glassy carbon electrode surface for physical adsorption of horseradish peroxidase [151]. Compared to electrochemically prepared polyaniline/polyvinylsulphonate films, the nanoPANI/dodecylbenzenesulphonic acid (DBSA) showed faster response time (0.62 s) and improved linearity for estimation of H2 O2 and better signal to noise ratio (61 ± 3). These results have been attributed to the highly ordered structure of nanoPANI/DBSA system. Jia et al. have fabricated horseradish peroxidase (HRP) biosensor by self-assembling gold nanoparticles to a thiol-containing sol–gel network of 3-mercaptaopropyltrimethoxysilane (MPS) [152]. The response time, detection limit, linear range and stability of the HRP biosensor for H2 O2 were found to be 2.5 s, 2.0 ␮mol L−1 , 5.0–10.0 ␮mol L−1 and 120 days, respectively. Ngmana et al. have immobilized horseradish peroxidase poly (2-methsulphonated polyaniline5-sulphonic acid)/l-lysine composite [153]. The correlation coefficient sensitivity, detection limit and the linear range of the amperometric HRP biosensor were determined to be 0.9966, 24.91 ␮A cm−2 , 0.01 mM, 0.01 mM of H2 O2 , respectively. Mathebe et al. have electrostatically immobilized horseradish peroxidase on the surface of polyaniline film electrochemically deposited onto platinum disc electrode [154]. The correlation coefficient and linear range of the polyaniline/peroxidase based biosensor were found to be 0.995 and 2.5 × 10−4 to 5 × 10−3 M, respectively. Zhou et al. have observed enhanced peroxidase activity in hemoglobin in sulfonated polyaniline on glassy carbon electrodes [155]. This result has been attributed to the facile interfacial transfer of hemoglobin mediated by sulphonated polyaniline. Lindgren et al. have immobilized horseradish peroxidase (HRP), sweet potato peroxidase (SPP) and peanut peroxidase (PNP) tobacco peroxidase (TOP) on graphite electrodes [156]. The peroxidase electrodes were used for estimation of H2 O2 . SPP was found to have the lowest detection limit. B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 Dai et al. electrochemically immobilized cytochrome C on a NaY Zeolite modified electrode [157]. This electrode displayed excellent response to the reduction of H2 O2 without the use of electron mediator and hence could be used for H2 O2 detection. Brahim et al. have developed a p (2-hydroxyethyl methacrylate/polypyrrole system containing glucose oxidase, cholesterol oxidase and galactose oxidase [158]. The observed stability of 9 months for this system indicates that this biosensor can be subcutaneously used to monitor glucose, cholesterol and galactose. Pijanowska et al. have made a comparison of urea estimation in blood plasma and in dialysate for the physiological and the pathological range by enzyme-field-effecttransistor (EnFET) based detection of pH and pNH4 [159]. It is concluded that pNH4 based biosensors require pre-dilution whereas for pH biosensors both for blood plasma and dialysate do not require dilution. Zhang et al. have discovered that stable mixed amphiphile(octadecylamine, ODA/behenic acid, BA)/urease Langmuir–Blodgett films can be formed onto the hydrolyzed surface of a pH-ISFET [160]. This urea biosensor has response time of 15 s, detection limit of 0.2 mM and linearity of 0–20 mM. Gambhir et al. have covalently immobilized urease on polypyrrole microspheres linked to conducting polypyrrole–polyvinylsulphonate (PPy–PVS) films [161]. These conducting polymer electrodes had a response time of about 40 s, were thermally stable for about 40 days at 25–50 ◦ C and could be used for estimation of urea from 5 × 10−3 to 6 × 10−2 mol L−1 using potentiometric technique for application to urea biosensor. Singhal et al. immobilized urease in mixed monolayers of polyvinylcarbazole (PNVK) and stearic acid (SA) [162]. Two values (9 and 30 mM) of Michaelis–Menton constant (Km ) were obtained at lower and higher concentrations of urea. These PNVK/SA/urease electrodes, stable for about 5 weeks at 4 ◦ C, were found to have detection limit and sensitivity as be 5 mM and 10 mV mM−1 , respectively. Rajesh et al. have fabricated an amperometric urea biosensor using covalent immobilization of urease onto poly (N-3-aminopropyl pyrroleco-pyrrole (PAPCP) copolymer film [163,242,244]. The copolymer electrode (PAPCP) was immersed in a phosphate buffer solution (0.1 M, pH 7.0) containing 0.015 M ethyl-dimethylaminopropylcarbodiimide (EDC) and 0.03 M and N-hydroxysuccinimide (NHS) for 1.5 h and was immediately placed in an enzyme solution of 10 mg mL−1 in the same buffer solution for another 1.5 h. These urease/PAPCP/ITO electrodes show stability for more than 2 months when stored at 4–6 ◦ C, linearity in the range of 0.16–5.02 mM, response time of 40 s, current sensitivity as 0.47 ␮A mM−1 cm−2 . The efforts are currently being made to improve the stability and ascertain the effect of interferents on these urea bioelectrodes. Rebriiev et al. have fabricated a urea biosensor based on ion-sensitive-field-effect-transistor (ISFET) and photopolymeric membrane obtained by irradiation of a liquid mixture comprising of vinylpyrrolidone, oligourethane methacryalate and oligocarbonate methacrylate [164]. This urea biosensor exhibited high sensitivity, response time as 20 min, linearity as 0.05–20 mM and stability for about 40 days. Sharma et al. have developed a monoenzyme amperomteric biosensor for the estimation of galactose in milk by immobi- 65 lizing galactose oxidase with poly (3-hexyllthiophene)/stearic acid onto indium–tin-oxide (ITO) coated glass plates using Langmuir–Blodgett film deposition technique [93–95]. The biosensor shows response time of 60 s, detection limit as 1 g dL−1 , linearity from 1 to 4 g dL−1 of galactose and shelflife as 90 days. These investigators later co-immbilized lactase and galactose oxidase onto Langmuir–Blodgett films of poly(3hexylthiophene)(P3HT)/stearic acid (SA) for estimation of lactose in milk and its products to prevent lactose intolerance [96]. These enzyme electrodes show a linearity of 1–6 g dL−1 of lactose, has a shelf-life of 120 days and could be used about ten times. These galactose oxidase (GaO) immobilized P3HT-SA electrodes were used for estimation of galactose in human blood [95]. The amperometric galactose biosensor based on P3HT had detection limit of 0.5 g L−1 of galactose and were stable for about 30 days. There was no interference with uric acid (0.1 mM), CaCl2 (0.1 mM) and ascorbic acid (0.2 mM). Pinto and Schanze have demonstrated the application of conjugated polyelectrolytes (CPEs) based on pendent ionic sulphonate and carboxylate groups for estimation of protease activity. The sensor has been used for estimation of low enzymeconcentrations and enzyme-catalyzed kinetics [168]. This technique can be used to develop assays for kinases, phosphates, lipases and esterases, etc. Ivanov et al. have used glassy carbon electrodes modified with polyaniline for the immobilization of cholinesterase using cross-linking technique [169]. Compared to other detectors, these polyaniline modified potentiometric biosensors utilized for the estimation of pesticides (trichlorfon, coumaphos, metiocarb and aldicarb) exhibit higher storage stability, increased sensitivities and lower detection limits. The degree of inhibition referred to as the total inhibitory effect of all the toxic compounds could be determined in about 15 min. Cosnier et al. have demonstrated that the immobilized thionine in the poly (dicarbazole-N-hydroxysuccinimide) results in the improved sensitivity (0.1457 A M−1 cm−2 ) and the maximum current (15.43 ␮A cm−2 ) of catechol of the polyphenol oxidase based biosensor [170,171]. These authors later fabricated an enzyme electrode by coupling of avidin and biotinylated polyphenol oxidase with biotinylated polycarbazole film resulting in improved biosensor stereoselectivity [170,171]. Rahman et al. have covalently immobilized pyruvate oxidase onto nanoparticles comprising of poly 5,2′ :5,2′′ -terthiophene3′ -craboxylic acid, poly-TTCA (nano-CP) on a glassy electrode for amperometric detection of phosphate ions [172]. The glassy carbon electrode (GCE) coated with nano-CP was immersed in a 20 mM HEPES buffer solution (pH 7.0) containing 10.0 mM ethyl-dimethyl-aminopropylcarbodiimide (EDC) for 6 h to activate the carboxylic groups of the nano-CP layers. The nanoCP/GCE was washed with HEPES buffer solution and subsequently incubated in 100 units of pyruvate oxidase in a 20 mM HEPES buffer (pH 7.0) solution for 12 h at 4 ◦ C. The immobilization of pyruvate oxidase occurred through the formation of covalent bond between carboxylic acid groups of the conducting polymer and amine groups of the enzyme. The detection limit, linearity, the response time and the shelf-life of this nanoCP biosensor was observed to be 0.3 ␮M, 1.0–100 M, 6 s and 1 month, respectively. 66 B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 Langer et al. have immobilized choline oxidase in nanostructured polyaniline layers of controlled nano and micro porosity for estimation of choline in food [173]. The choline oxidase based polyaniline biosensor is stable for about 30 days, has sensitivity of 5–10 mV mM−1 in the potentiometric mode and 5 ␮A mM−1 in the amperometric mode. Qu et al. have reported an amperomteric biosensor for detection of choline based on polyaniline multilayer film and layer-by-layer assembled functionalized carbon nanotube [174]. The choline biosensor showed response time as 3 s, detection limit as 0.3 ␮M and linear range as 1 × 10−6 to 2 × 10−3 ␮M. Biloivan et al. demonstrated that protein selective microelectrodes can be fabricated by immobilization of different proteinases such as trypsin, pronase E and carboxypeptidase B on surface of platinum electrodes modified by conducting PPy[3,3Co(1,2-C2 B9 H11 )]2 . This biosensor had a detection limit of 1–2 ␮g mL−1 in phosphate buffer [175]. 3.2. Conducting polymer based microbial biosensors Microorganisms as sensing elements provide advantages like over-coming the process of purification for enzymes, multi-enzyme functions and cofactor/coenzyme addition, etc. [176–180]. The recent progress in molecular biology/recombinant DNA technologies has opened enormous possibilities of tailoring microorganisms to improve the activity of an existing enzyme or express foreign enzyme/protein in a host cell such that it can be utilized to enhance specific activity. There have been various strategies to modify the microbes for application to microbial biosensors such as lux gene expressions for fluorescent protein, gfp gene for green fluorescence protein. The enhanced expressions of gene of interest have been reported to improve the sensitivity for target analytes. Besides this, various electrochemical detection methods have been used to integrate microbes with various immobilization matrices. Kwok et al. immobilized Bacillus subtilis in a sol–gel derived biocompatible material of silica and poly (vinyl alcohol)-graftedpoly (4-vinylpyridine) (PVA-g-PVP) copolymer resulting in fast diffusion of organic compounds [177]. This prototype optical biosensor can be used for the estimation of biochemical oxygen demand (BOD) in waste water samples. Lei et al. have reported a whole cell biosensor using PNP-degrader P. putida JS444 for estimation of organo-phosphorus nerve agent with p-nitrophenyl constituent [178]. This biosensor was found to be stable for only 5 days. Applications such as BOD sensing, toxicity sensing, microbial fuel cell and various other important biological parameters such as alcohol, sugar, phenol, methanol, heavy metals, cyanide, etc. appear to be the most promising. Efforts are being made to commercialize some of microbial sensors through optimization of the microbial species, immobilization medium, oxygen sensing film and suitable engineering design. Minet et al. have reported that a mediated (toluidine blue) antibody (Salmonella and E. Coli)-immobilized polypyrrole electrode can be used for detection of Listeria monocytogenes using amperometric method [180]. This biosensor could reproducibly and selectively detect Listeria after incubation in the micro-organism at levels of 105 cells mL−1 in 30 min. 3.3. Conducting polymer based immunosensors Immunosensors can be used to detect the presence of an analyte in minutes using small volumes of sample [181–190]. An effective combination of immunochemistry coupled with electro-chemistry in an analytical device could provide the basis of direct electrical detection for a wide range of analytes with specificity and great sensitivity. A number of immunosensors based on conducting polymers have been reported. A unique approach for conducting polymers reversible immunosensors using pulsed electro-chemical detection has been developed. The strategy involves the detection of organochlorine pesticides including PCBs, atrazines and chlorinated phenols. Porter has investigated electro-polymerized conducting polymers as antibody receptors in [181]. They have shown that antibodies against conducting polymers (carbazole as a hapten) may react to modulate the polymer electrochemistry. The reaction of the anti-serum was found to influence the polymer electrochemistry by amperometric response and therefore can be utilized as a sensor for amino acid. Grant et al. have reported a label free and reagentless immunosensor based on direct incorporation of antibodies into conducting polymer films using ac impedometric electrochemical interrogation [182]. It has been demonstrated that faradic real component of the impedometric response acts as a dominant component of AC impedometric response of anti BSA loaded conducting polypyrrole (PPy) film on its exposure to different concentration of BSA. The nature of observed Faradaic response current, perhaps arising due to antibody–antigen interaction in this system is still a matter of speculation. Ramanaviciene and Ramanavicius have discussed the use of conducting polymers thin films for application as electrochemical affinity sensors with the emphasis on design and applications of novel immunosensors [184]. They have briefly discussed the biological active component for the creation of polypyrrole based immunosensors. Gooding et al. have fabricated the glassy carbon electrode modified with anti-rabbit IgG antibody entrapped in an electrodeposited polypyrrole membrane for label free amperometric detection of rabbit IgG antigen in flow injection system [183]. The observed reversible antibody–antigen binding was assigned to the short anodic pulse (200 ms). The changes in the concentration of the immobilized antibody within a polymer matrix significantly affected the stability of the immuno-electrode. Zhang et al. have fabricated a low-cost label-free amperometric immunosensor based on anti-rubella serum immobilized onto nano-Au/poly-o-phenyldiamine doped with Prussian blue for the detection of rubella vaccine [185]. This immunosensor had a signal to noise ratio of 3 and exhibited response to rubella vaccine in the range of 8.1 × 10−8 to 8.0 × 10−6 cell culture infective dose (CCDI) mL−1 . Tahir et al. have described the characteristics of polyaniline compounds in different protonic acid for application to diarrhea virus detection based on polyclonal and monoclonal BVDV antibodies [186]. They have shown that PANI with perchloric acid show highest conductivity in pH 6.6 and the sensitivity of the biosensor ranged from 103 to 104 cell culture infective dose (CCDI) mL−1 . They however B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 did not investigate the effect of polymerization conditions on the characteristics of the biosensor. Darain et al. have developed a mediator-less immunosensor for the detection of rabbit IgG (RIgG) by covalently bonding horseradish peroxidase and streptavidin onto conducting polymer(5′ ,2:5′ 2′′ -terthiophene-3′ -carboxylic acid) coated on a screen-printed electrode [188]. The catalytic current measured amperometrically at −0.35 V versus Ag/AgCl showed detection limit as 0.33 ␮g mL−1 and a linear range of RIgG concentration from 0.5 to 2 ␮g mL−1 . Lillie et al. fabricated simple immunosensor formats by polymerizing pyrrole loaded with avidin or antibody to luteinising harmone (LH) on a gold inter-digited electrode and demonstrated that impedance spectroscopy can be used to detect LH between 1 and 800 IU L−1 [189]. Farace et al. have developed a reagentless immunosensor for detection of luteinising harmone based on antibody entrapped in a conducting polypyrrole matrix using impedance spectroscopy [190]. Interestingly, the interaction of DNA with conducting polypyrrole could be used to differentiate single- and double-stranded DNA. 3.4. Conducting polymer based nucleic acid biosensors DNA probes and biosensors have widely attracted much attention for the diagnosis of various disorders [191–219]. DNA biosensors have enormous applications in clinical diagnostics for inherited diseases, rapid detection of pathogenic infections, and screening of c-DNA colonies required in molecular biology. Conventional methods for the analysis of specific gene sequences are based on either direct sequencing or DNA hybridization. Lucarelli et al. have recently reviewed the literature in relation to the application of gold and carbon electrodes as electrochemical transducers for DNA hybridization sensors [202]. Because of its simplicity, most traditional techniques in molecular biology are based on hybridization. Marrazza et al. have reported an electrochemical DNA biosensor based on covalent immobilization of double stranded calf thymus DNA on the surface of a disposable screen-printed graphite electrode using avidin–biotin coupling [187]. Differential pulse voltammetry (DPV) and potentiometric stripping analysis at constant current were used to detect a specific sequence of bases and the presence of a gene or an infectious agent of the specific oligonucleotide target in solution to the concentration of 2 ␮g mL−1 necessitating amplification of hybridization signal for real samples. In this context, conducting polymers are reported to be good candidates for DNA immobilization [161]. Gambhir et al. have immobilized calf thymus DNA on electrochemically prepared conducting polypyrrole/polyvinyl sulphonate films [193]. Ban et al. have developed a method for investigating DNA–protein interaction using electrochemical techniques [210]. Yu et al. have developed an amperometric enzyme-linked immunoassays based on single-wall carbon nanotubes (SWNT forests) on pyrolytic graphite surfaces for estimation of human serum albumin (HSA) using horseradish peroxidase (HRP) labels [194]. Using mediator (hydroquinone) alongwith SWNT provided the detection limit of HSA to 1 pmol mL−1 (1 nM). Some of the issues currently being addressed relate to the pat- 67 terning of SWNT forests into arrays for multiple protein assays and the strategy to avoid soluble mediation. Cha et al. have reported the application of conducting poly (thiophen-3-yl-acetic acid 1,3-dioxo-1,3-dihydro-isoindol-2-ylester) (PTAE) to DNA hybridization electrochemical sensor [199]. This sensor has a sensitivity of 0.62 ␮A nmol−1 and detection limit as 1 nmole of target oligonucleotides (ODN). The limited selectivity of this sensor has been assigned to the increased steric hindrances resulting due to partial destruction of double helix structure. The physical and chemical conditions affecting the response have not yet been studied. Youssufi and Makrouf have prepared conducting polypyrrole substituted with ferrocenyl groups as DNA electrochemical sensor [203]. This sensor has a detection limit of less than 1 pmol of DNA target. Both the thermal stability and the effect of interferents on the response of the DNA biosensor have not yet been reported. Gu and Hasebe have incorporated horseradish peroxidase (HRP) and methylene blue (MB) in polyion complex membrane comprising of DNA (ds-DNA) and poly (allylamine) (PAA) on the surface of gold (Au) disk electrode to fabricate second generation H2 O2 biosensor [204]. This HRP/DNA-MB/PAA/Au bioelecrode had high sensitivity (14.6 ␮A−1 mM−1 ) and detection limit as 0.3 ␮M. The attempts are being made to apply this electrode to other oxidases that produce H2 O2 . Shi et al. have coated poly (anilineaniline boronic acid) on DNA templates to fabricate redox-active polymer wires that electrically contact glucose oxidase reconstituted on the polymer wires with the electrode [205]. It is envisaged that glucose oxidase may be similarly contacted with negatively charged polyelectrolyte templates like polystyrene sulphonate. Arora et al. have physically immobilized double stranded calf thymus deoxyribonucleic acid (DNA) onto electrochemically prepared polypyrrole-polyvinylsulphonate (PPyPVS) films [90]. The amperometric response studies of the DNA/PPY-PVS electrodes carried out at 25 ◦ C as a function of 2-aminoanthracene (2 AA) concentration (0.01–20 ppm) and o-chlorophenol (OCP, 0.1–30 ppm) reveal that 10 ppm is sufficient to reduce the observed guanine oxidation current and that 25 ppm of OCP reduced the oxidation current to zero. The effect of interferents like calcium, magnesium carbaryl and bisphenols on the response of the DNA/PPY-PVS electrodes has not yet been investigated. Wu et al. have reported a biosensor based on polyaniline intercalated graphite oxide nanocomposite (PAI/GO) for monitoring DNA hybridization [206]. These workers observed that these electrodes without ss-DNA exhibited linear behaviour between 34 and 241 ␮g mL−1 of ss-DNA. However, when the PAI/GO electrode was immobilized with ss-DNA, the hybridization current was found to be linear from 275 to 241 ␮g mL−1 of complementary ss-DNA. These results showed that paraffin–graphite–PAI/GO electrode can be used for the estimation of ss-DNA. The effect of temperature and the various interferents on the response characteristics of the modified PAI/GO electrode has not yet been studied. Kerman et al. have fabricated a magnetic bead-based hybridization assay for application as peptide nucleic acid probes. This method can be used to detect specific sequences in PCR amplified DNA samples [211]. Fan et al. have fabricated an electrochemical DNA sensor based on molecular beacon 68 Table 1 Characteristics of various conducting polymer based biosensors S. no. Bioensing molecule Analyte Immobilization method Method of detection Response time Detection limit Linearity Interference Stability Reference Polypyrrole Virostate anti-AZT AZT 10 pg mL−1 – – – [58] Poly(oaminophenol/carbon nanotubes) Ferocenyl carbon nanotubes/polypyrrole Polysiloxane/polypyrrole Glucose oxidase Glucose Electrochemicalimpedance Amperometric – 2 Electrochemical entrapment Electro-Polymerization 3s 0.01 mm Upto 5 mm 30 days [118] Glucose oxidase Glucose Electro-polymerization Amperomteric – – – No interference from ascorbic acid, uric acid and acetaminophen – – [122] Tyrosinase Catechol Electrochemical entrapment Spectro-photometric – – – – 35 days [128] 5 Polypyrrole-poly(vinyl alcohol) Methanol Resistance change 5 min approx – 50–1059 ppm – 2 weeks [130] 6 Polyaniline Polypyrrolepoly(vinyl alcohol) film Uricase 60 days [134] 7 No interference from ascorbic acid – – [149] 3 4 8 9 10 11 12 13 Uric acid Template Amperometric 10 s – 0.0036–1 mmol dm3 HRP Poly(3,4ethelenedioxythiophene)poly(styrenesulfonate) + poly4-vinylpyridine Lactate oxidase Poly(5-hydroxy-1,4naphthoquinone-co-5hydroxy-3-acitic acid-1,4-naphthoquinone) Polyaniline/polyvinylsulphonate Alcohol dehydrogenase Octadecylamine Urease H2 O2 Entrapment – – 0–30 ␮M – Lactate Covalent Amperometric – 50 ␮M 50–1500 ␮M No interference from ascorbic acid, glycine and acetaminophen 9 weeks [142] ␤-nicotinamide adenine dinucleotide Urea Solution phase Amperometric – – – – – [143] Potentiometric 15 s 0.2 mm 0–20 mm – – [160] Polyanilinedodecylbenzenesulphonic Poly(phenyleneethynyleneSO3 /CO2 Glassy carbon-polyaniline HRP H2 O2 Adsorption in the subhase of Langmuir Layer Electrostatic adsorption Colorimetric – – – – – [151] Protease P-Nitranilide and rhodamine-Arg-2 Organophosphorus and carbamic pesticides BOD in waste water – 3 min – – – – [168] Cross-linking Fluorosscence spectroscopy – – Coumaphos – – – [169] – Optical 20 min – Heavy metals (transition metals) 45 days [177] 0.2–50 ␮M−1 5 days [178] 0–75 ppm No interference from phenolic compounds and pesticides – – [182] – [183] 60 days [185] – [187] Cholinesterase 5 × 10−9 Oxygen sensing film 15 Silica-tris(4,7-diphenyl1,10-phenanthroline) ruthenium (II) Oxygen sensing film Pseudomonas putida JS444 Organophosphates – O2 consumption Less than 5 min 25 mg L−1 for B. subtilis and 60 mg L−1 for activated sludge Lower than 60 ppb 16 Polypyrrole Anti-BSA BSA ac Impedance – – 17 Glassy carbon/polypyrrole Anti-rabbit IgG Rabbit IgG Electrochemical entrapment Entrapment Amperometric – – 8.1 × 10−8 14 18 Nano Au-ophenylenediamine/prussian blue Anti-rubella Rubella Physical Amperometric 6 min 4.010–8 lgccid 50 mL−1 19 Graphite screen printed electrode DNA DNA Biotin–avidin/adsorption – 1 ␮g mL−1 – 20 Polypyrrole-polyvinyl sulphonate Single wall carbon Nanotube PTAE Ferrocenyl popypyrrole DNA OCP Physical Differential pulse voltammetry and chronopotetiometry Amperometric No interference from human serum No significant interference rom Mumps, varicella, human serum albumin, etc. – 30 s – 0.1–25 ppm – 4 months [90] Anti-human serum albumin (HSA) DNA Amino-DNA Human serum albumin (HSA) DNA DNA Covalent Amperometric 1 nm – – – [193] Substitution Covalent Cyclic voltammetry Cyclic voltammetry 1 nmole 10−14 mol – – – – – – [199] [203] 21 22 23 to 8 × 10−6 lgccid 50 mL−1 B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 Matrix 1 24 25 PNA/polymer/nuclease enzyme DNA–poly(allylamine) Ferrocene tagged PNA HRP/methylene blue 26 Polyaniline 27 Polyaniline–graphite oxide nano composite in carbon paste electrode Poly(9,9-bis(6′ -N,N,Ntrimethylammoniumhexyl)fluorene phenyle) Polyttca/glassycarbon Polyttca/carbon Cholesterol oxidase, cholesterol esterase and peroxidase DNA 28 29 30 Fluoroscence 2–3 s for each addition <10 s – – – – [205] 0.3 ␮M Upto 0.1 mm 5 days [206] 6 weeks [249] H 2 O2 Solution cast Square wave voltammetry Cholesterololeate and cholesterol Covalent Spectroscopic and amperometric 240 s 25 mg dL−1 25–500 mg dL−1 No interference from ascorbic acid, l-cysteine and uric acid – DNA Physical adsorption Conductometry – – – – – [208] Peptide nucleic acid DNA Solution phase Fluorescence detection – 10 pm – – – [209] Pyruvate oxidase HRP/streptavidin/Anti rabit igg Polyphenol oxidase Phosphate ions Rabit igg Covalent immobilization Avidin–biotin coupling Amperometric Amperometric 6s 35 min 0.3 ␮M 0.33 ␮g m L−1 1.0–100 ␮M 0.5–2 ␮g mL−1 No interference – 1 month – [172] [188] Thionine Grafting Amperometric – – – – – [170] HRP H 2 O2 Drop coating Amperometric – 0.01 mm 0.01–0.1 mm No interference of ascorbic acid 6 months [153] Glucose oxidase Glucose Galvanostatic – – 0–10 mm 14 days [123] – – 0–35 mm 1 month [173] 200 s 0.3 mm 0.1–2 mm – [76] 1 month [174] 33 Poly(dicarbazole-Nhydroxysuccinim ide) Poly(2-methoxyaniline-5sulfonic acid) Polypyrrole 34 Polyaniline Choline oxidase Choline Galvanostatic electropolymerization Diffuion 35 PEDOT MIP Morphine Precipitation Amperometric and potentiometric Amperometric 36 Polyanilne/multi-wall nanotube (MWNT) Polyanilie Poly 3,4-ethylenedioxythi ophene Polypyrrole Polypyrrole Thiophene capped polytetrahydrofuran (TPTHF)-co-pyrrole (Py) 1,2-Diaminobenzene Polypyrrole Choline oxidase Choline Cross linking Amperometric 3s 0.3 ␮M 1 × 10−6 to 2 × 10−3 M No interference of ascorbic acid No interference of codeine No interfere nce HRP Tyrosinase H2 O2 Phenolic compounds Electrostatic Entrapment Amperometric Amperometric – 20–40 s 2.5 × 10−4 M – 2.5 × 10−4 to 5 × 10−3 M 5–500 nm – – – 12 days [154] [140] HRP Anti-listeria Polyphenol oxidase H2 O2 Listeria Phenolic compounds Entrapment Entrapment/covalent Entrapment Amperometric Cyclic voltammetry Spectrophotometric method <20 s 30 min – 5 × 10−5 105 cell mL−1 – 9.0 × 10−7 –2.0 × 10−4 M – – – – – 2 weeks – 20 days [131] [180] [141] Glucose oxidase Trypsin Covalent Cross-linking Amperometrically Potentiometric 15 s – 0.05 mmol L−1 0.05 mmol 0.05–5 mmol L−1 0.05–0.5 mmol No interference – 90 days – [119] [175] DNA Entrapment ac Impedance 30 min – 1.0 × 10−5 –3.0 × 10−8 mol L−1– – [219] 45 46 Polypyrrole–MWNTCOOH Polypyrrole Polypyrrole Glucose N␣-benzoyl-l-arginine ethylester hydrochlori de DNA DNA DNA DNA DNA Entrapment Covalent – – – – – – – – – – [217] [192] 47 Polyterthiophene DNA DNA Covalent – – – – – [218] 48 Polyaniline DNA DNA Covalent Microgravimetric Photocurrent spectroscopy Impedance and admittance Differential pulse voltammetry – 1.0 × 10−12 mol L−1 – – – [212] 32 37 38 39 40 41 42 43 44 B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 31 Snps 69 70 B.D. Malhotra et al. / Analytica Chimica Acta 578 (2006) 59–74 like DNA stem–loop labeled with an electro-active reporter as the hybridization sensing element [207]. Such an E-DNA sensor can be used to electrochemically detect a target-induced conformation change in a biopolymer. The synthetic DNA [peptide nucleic acid (PNA)] due to its superior enzymatic stability and hybridization characteristics coupled with unique biochemical and physico-chemical properties probe has been used to estimate DNA target upto about 10 femto moles (fM) using conductance measurements. Gaylord et al. have demonstrated that fluorescently labeled PNA coupled to conducting poly (9,9bis (6′ -N,N,N-trimethylammoniumhexylbromide)fluorine)co-phenylene and a nuclease can be used to identify single nucleotide polymorphism (SNP) [209]. The addition of an enzyme prior to the addition of a conjugated polymer has been found to increase the DNA target lengths and the selectivity of the assay. Besides this, conducting polymer/PNA system provides added sensitivity to standard single fluorophore-labeled probes without the need for complex/expensive optical setups for identification of neurodegenerative diseases. It may be remarked that PNA is uniquely placed compared to many other nucleic acid derivatives. Analysis using bead-probe techniques and light-up probes reveal the degree of improvement that can be achieved. In spite of the high cost of the synthesis, PNA provides a unique opportunity for advanced high biosensor technology with regards to sensitivity, stability, automation and system integration. Zhu et al. have developed a novel and sensitive electrochemical DNA biosensor based on electrochemically fabricated polyaniline nanowire and methylene blue for detection of DNA hybridization [212]. The detection limit of the DNA biosensor for discrimination of complementary and non-complementary DNA sequences was found to be 1 × 10−12 mol L−1 . Taira et al. have immobilized single stranded DNA probe on self-assembling polymer based on polyallyamine modified with thioctic acid for hybridization assays [213]. The efficiency and selectivity of hybridization assay could be affected by adjusting the ionic strength of sodium chloride. Ioannou et al. have reported an electrochemical DNA biosensor based on carbon paste electrodes modified by a conducting composite for investigation of DNA-drug interactions [214]. Lori et al. developed a conducting polymer surface by doping N-nitriloacetic acid (NTA) into the electropolymerized polyvinylsulphonate doped polyaniline (PANI/PVS) at a screen printed carbon electrode for the immobilization of his-tagged biomolecules [215]. The resulting NTA-PANI-/PVS film was shown to have interesting electrochemical properties. Crucifix et al. have investigated the immobilization of double stranded DNA molecules and of nucleoprotein complexes on 2D-strepavidin crystals [216]. The binding efficiency and specificity were examined using radioactively labeled oligonucleotides by direct visualization of unstained and hydrated nucleic acid molecules using cryo-electron microscopy. It is shown that RNA polymerase, once immobilized, femtomolar amounts of DNA template can suitably interact with the enzyme. Lassalle et al. have investigated the electrosynthesis of conducting copolymer using pyrrole and pyrrole-oligonucleotide (ODN) at platinum disc electrodes [192]. The poly (pyrrole/ pyrrole-ODN) films could be used to detect DNA hybridization in real time using quartz crystal microbalance (QCM). The affinity between avidin and biotin modified DNA target was used to validate the effectiveness of transduction by fluorescence microscopy [217]. Lee et al. have developed an electrochemical method to directly detect DNA hybridization using a new conducting polymer, which was polymerized on the glassy carbon electrode with a terthiophene monomer having a carboxyl group (3′ -carboxyl-5,2′ ,5′ ,2′′ -terthiophene) [218]. The highest difference in admittance was seen at 1 kHz before and after hybridization. Peng et al. have reported a novel oligonucleotide (ODN) sensor based on electropolymerization of polypyrrole in the presence of a sample containing ODN(s) [240]. The resulting trapped ODN(s) have been probed by addition of complementary sequence ODN using impedance spectroscopy. The impedance signal was found to be almost linear in the range of 3.7–370 nM with detection limit of ∼1 nM. By incorporating cadmium sulphide (CdS) nanoparticles with the probe, a significant improvement in sensitivity was obtained. Cai et al. have demonstrated that electrochemical impedance spectroscopy can be used to visualize charge transfer through conducting polypyrrole films loaded with oligonucleotides probes formed on the carbon nanotube modified electrodes as a basis for reagentless protocol [219]. This technique has several advantages such as high selectivity, reduced reaction time without the use of mediators or fluorescent materials for complementary and mismatched target sequences. Table 1 gives the characteristics of the various conducting polymer based biosensors discussed in the above sections. 4. Conclusions An attempt has been made to present an overview on the prospects of conducting polymers in biosensors reported since 2000 till date. The extended ␲ systems in conjugated polymers, highly susceptible to oxidation or reduction, provide precise control to the electrical and optical properties as biochemical reactions are often reversible in nature. 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