Signal Generation at an Electrochemical Immunosensor via the Direct Oxidation of an Electroactive Label

Full Paper

Signal Generation at an Electrochemical Immunosensor via theDirect Oxidation of an Electroactive Label

Mohammad Akram,a Margaret C.Stuart,b Danny K. Y. Wonga*a Department of Chemistry, Macquarie University, Sydney, New South Wales 2109, Australiab Department of Biological Sciences, Macquarie University, Sydney, New South Wales 2109, Australia*e-mail: [email protected]

Received: August 26, 2005Accepted: November 7, 2005

AbstractIn this paper, we report the development of an electrochemical immunosensor with a direct signal generation scheme.To achieve this, ferrocenecarboxylic acid was conjugated as an electroactive label to an immunocomplex immobilizedon an electrochemically pretreated glassy carbon electrode. The pregnancy hormone and tumor marker, humanchorionic gonadotrophin (hCG), was employed as a model analyte. Direct oxidation of ferrocenecarboxylic acidfacilitated a signal generation scheme for the quantification of hCG at the immunosensor. Evaluation of theimmunosensor performance has yielded statistically acceptable agreement in hCG levels present in several humanserum samples compared to those reported by a Sydney hospital.

Keywords: Electrochemical immunosensors, hCG detection, Ferrocenecarboxylic acid, Direct signal-generationscheme, Non-competitive immunoassays

DOI: 10.1002/elan.200503402

1. Introduction

There is a continuing demand for fast and simple analyticalmethods for thedeterminationofmany clinical, biochemicaland environmental analytes. Immunosensors based on theuse of antibodies as specific binding reagents provide apromising means of analysis owing to their specificity andsensitivity. Specificity arises from themolecular recognitionof target analytes by antibodies to form a stable complex onthe immunosensor surface [1, 2]. On the other hand,sensitivity depends on several factors including the use ofhigh affinity analyte-specific antibody(ies), their orientationafter being immobilized on the immunosensor surface andthe appropriate detection system for measurement of thesignal generated via a transducer [3]. Many types ofimmunosensors have been developed according to themode of detection. Examples of these include electro-chemical [4 – 6], chemiluminescent [7], fluorometric [8],piezoelectric [9] and surface plasmon resonance immuno-sensors [10], impedance biosensing chip [11] and multi-channel microchips for multianalyte determination [12].In developing an electrochemical immunosensor, as most

antibodies and antigens are intrinsically unable to act asredox partners, an enzyme label is often conjugated to aparticular component of the immunocomplex to promote anelectrochemical reaction. This is achieved by enzymaticconversion of the substrate into an electroactive product,which subsequently undergoes a redox reaction uponapplying a suitable potential. The reduction or oxidationcurrent can thenbemeasured for quantification purposes. In

such a signal-generating scheme, the catalytic reactions ofenzymes enable an amplification of the signal that can bedetected even when a minute quantity of enzyme is used.Alkaline phosphatase and horseradish peroxidase are twocommon enzymes conjugated to antibodies for signalgeneration in the development of specific and sensitiveimmunosensors by a number of investigators [13 – 18].Notably, an enzymatic signal-generating scheme is anindirect detection technique. Moreover, alkaline phospha-tase, which is present in biological samples such as serumand plasma, could interfere in assays if it were not removedby washing steps following specific interaction of sampleanalyte with its specific antibody reagent. Further, as somehorseradish peroxidase substrates are carcinogenic [19], it isimportant to select one that is safe for routineoperationwithan immunosensor.There has previously been an interest in developing a

more direct signal generating technique to facilitate simul-taneous immunoassays. For example, Hayes et al. [20]demonstrated the anodic stripping voltammetric detectionof Bi3þ and In3þ after these metal ions were released fromBi3þ and In3þ – labeled-proteins. The stripping current wasthen used as ameasure of the amount of proteins of interest.In a different study, Yasuzawa et al. [21] employed ferro-cene as an electroactive label conjugated to the signalantibody of the analyte transferrin. The magnitude of theoxidation current of ferrocene was then quantitativelyrelated to the amount of transferrin present. Similarly,Lim and Matsunaga [22] reported the preparation of aferrocenecarboxylic acid-conjugated antibody for use as an

237

Electroanalysis 18, 2006, No. 3, 237 – 246 D 2006 WILEY-VCH Verlag GmbH&Co. KGaA, Weinheim

analytical reagent in the separation of immune complexesfrom free conjugate in a cation exchange capillary column,based on their different isoelectric points. In this way, thecurrent produced by the oxidation of ferrocenecarboxylicacid was measured to determine the quantity of conjugatedimmunocomplex.More recently,Okochi et al. [23] have alsoemployed antibodies conjugated with ferrocencarboalde-hyde as a probe for the detection of immunoglobulinG in anon-chip type flow immunoassay. In general, ferrocenederivatives are known to be fast, reversible redox com-pounds that can be used as mediators in enzyme biosensors[24 – 27]. Among them, ferrocenecarboxylic acid is useful asit is a water-soluble derivative and can be easily conjugatedto immunoglobulin G antibody via 1-ethyl-3-(3-dimethyla-minopropyl) carbodiimide hydrochloride (EDC)-modifiedcarboxylic acid terminals [28]. Upon applying an appropri-ate potential, the ferrocenecarboxylic acid is oxidizedrapidly without involving any intermediate product. There-fore, ferrocenecarboxylic acid is a superior candidate for useas an electroactive label conjugated to a biological compo-nent for development of a direct signal-generating schemeat an electrochemical immunosensor.In this paper, we report the incorporation of ferrocene-

carboxylic acid in a direct and rapid signal-generationscheme of a molecular architecture immobilized on anelectrochemical immunosensor. In our design, the ferroce-necarboxylic acid-conjugated antibodies are notably posi-tioned closer to the electrode surface than that involvinglabeled antibodies in an aqueous solution. This aids inreducing diffusion of ferrocenecarboxylic acid towards theelectrode surface prior to oxidation, and hence contributesto the sensitivity level achievable by an electrochemicalimmunosensor. The tumor marker, human chorionic gona-dotrophin, has been used as amodel analyte to demonstratethe application of this immunosensor for analysis of clinicalsamples.

2. Experimental

2.1. Materials

Two murine monoclonal antibodies specific for humanchorionic gonadotrophin (hCG), Ab1 (clone 26/1) and Ab2(clone 89/9), in 0.05 M Tris acetate (pH 7.5) with 0.05%Bronidox, were purchased from Bioclone Pty Ltd. (NSW,Australia). Iodination grade hCG was purchased fromMerck Pty Ltd (Vic, Australia). Goat anti-mouse anti-serum for radioimmunoassay was purchased from Chem-icon Australia Pty Ltd (Vic, Australia). 125I-iodine used forlabeling Ab1 and hCG was purchased as Na125I in sodiumhydroxide from Australian Nuclear Science & TechnologyOrganization (NSW, Australia). Chloramine-T and sodiummeta-bisulfite (both from BDH, NSW, Australia) were usedas received. Ferrocenecarboxylic acid, EDC, bovine serumalbumin (BSA), sodium 4-(2-hydroxyethyl)-1-piperazinee-thanasulfonate (HEPES) and polyethylene glycol (PEG)6000 were purchased from SigmaAldrich (NSW,Australia).

Reagent grade KNO3, Tris, MgCl2, NaCl, KH2PO4, EDTA(disodium salt), and Na2HPO4 (all purchased from AjaxChemicals, NSW, Australia) were used for the preparationof phosphate-saline buffer (10 mM phosphate and 150 mMNaCl, 10 mM EDTA) (PBS, pH 7.6). Alumina powder ofmesh size 1.0, 0.3 and 0.05 mm was purchased from LecoCorporation (MI, USA). Aqueous solutions were preparedusing Milli-Q water (Milli-Q Reagent Water System). A0.5% (w/v) BSA-PBS buffer (pH 7.6) was used, whererequired, throughout the experiments. In all voltammetricexperiments, analyte solutions were purged with nitrogenfor 10 min. After deaeration, a blanket of nitrogen was keptover the solution. A further 1 min was allowed for thesolution to become quiescent prior to the commencement ofeach experiment.

2.2. Instrumentation

A 10-mL electrochemical cell accommodated with a three-electrode system consisting of an immunoelectrode (seebelow), a Ag jAgCl reference electrode (BioanalyticalSystems Inc., IN, USA) and a platinum coil as auxiliaryelectrode, was used throughout. All electrochemical experi-ments were performed using a MacLab/4e potentiostat(ADInstruments Pty Ltd, NSW,Australia) interfacedwith aMacintosh computer (with OS-9 operating system), usingv1.3 – 1 EChem and v4.2Chart software (ADInstrumentsPty Ltd). Microcon centrifugal filtration devices with a10,000-Dalton molecular weight cut off were purchasedfrom Millipore Pty Ltd. (NSW, Australia). Either an LKB-Wallac 1260 Multigamma II counter or a RackGammacounter (LKBAustralia Pty Ltd, NSW, Australia) was usedfor radioactivity measurements. PD-10 columns containinga Sephadex G-25 matrix (Calbiochem, NSW, Australia)were used in all gel filtration experiments.

2.3. Labeling of Ab1 and hCG with 125I

AChloramine-T method was employed to label 7.5 mg and10 mg of Ab1 and hCG, respectively, with approximately 37MBq of 125I-iodine according to the method of Greenwoodand Hunter [29]. Briefly, 10 mL of Na125I were added to100 mL of 100 mM phosphate containing the protein ofinterest. Following the addition of 5 mL of 1 mg mL�1

aqueous Chloramine-T, the reaction proceeded for 40 s ona vortex mixer before quenching by the addition of 30 mL of2 mg mL�1 sodiummeta-bisulfite. Separation of 125I-Ab1 or125I-hCG from unreacted 125I was performed by gel filtrationon a PD-10 column. Prior to transferring the reactionmixture to the column, 3-bed volumes (each 9 mL) of 0.5%BSA-PBS buffer were passed through the column. Thereaction mixture was applied to the column and eluted with0.5% BSA-PBS buffer. Thirty 500 mL fractions were col-lected, and the radioactivity of a 5 mL aliquot of eachfraction was counted using a gamma counter. The eluted

238 M. Akram et al.

Electroanalysis 18, 2006, No. 3, 237 – 246 www.electroanalysis.wiley-vch.de D 2006 WILEY-VCH Verlag GmbH&Co. KGaA, Weinheim

www.electroanalysis.wiley-vch.de

fractions containing 125I-Ab1 or 125I-hCG were dispensedinto 10 mL and 50 mL aliquots and stored at�20 8Cuntil use.

2.4. Preparation of Ab2 Ferrocenecarboxylic AcidConjugates

Conjugation of ferrocenecarboxylic acid to hCG antibodywas carried out according to the procedure describedbyLimand Matsunaga [22]. Briefly, 4 mg of ferrocenecarboxylicacid were dissolved in 800 mL of 0.15 M HEPES buffer(pH 7.3) and the solution was filtered through a 0.2-mmpore-diameter membrane. Then, 10 mgEDCwere added tothe solution and mixed until the solid was completelydissolved and the solution was homogeneous before adding90 mL of Ab2 (500 mg mL�1 in PBS). The mixture wasincubated for 2 h at room temperature on a shaker. Afterincubation, the sample was centrifuged at 14,000 g for30 min in a Microcon filter at 4 8C. The residue in themicrocon filter containing conjugated antibody was washedwith 4 or 5 volumes of 200 mL PBS to ensure that uncon-jugated ferrocenecarboxylic acid was removed. The amountof iron present in the ferrocenecarboxylic acid-antibodycomplex was determined by measuring the absorbance at562 nm [30]. The results showed that no iron was detectablein the antibody preparation prior to conjugation but therewere 13 iron atoms per antibodymolecule after conjugation.This result is consistent with that reported by Lim andMatsunaga [22]. The antibody content of the conjugate wasdetermined by a competitive radioimmunoassay procedure.Dilutions of ferrocenecarboxylic acid-Ab2 conjugate andstock Ab2 were incubated with 125I-labelled hCG overnightat room temperature in a final volume of 400 mL 0.5%BSA-PBS. Next day, hCG– antibody complexes were precipitat-ed by the addition of normal mouse serum, goat anti-mouseserum and PEG 6000 at a final concentration of 6%. Aftercentrifugation, the supernatant was decanted and the radio-activity in the precipitates was counted. The antibodycontent of the ferrocenecaboxylic acid-Ab2 complex wasdetermined by reference to a stock Ab2 dilution curve.Conjugated antibodywas reported to be homogeneous [22].This conjugate was used in all subsequent experiments.

2.5. Immunoelectrode Preparation

Before use, bare 3-mm diameter glassy carbon electrodes(Bioanalytical Systems Inc., IN, USA) were polishedsequentially in slurry of alumina powder of 1.0, 0.3 and0.05 mm to a mirror finish. The electrodes were thensonicated in Milli-Q water for 2 min to remove any aluminaremnants and rinsed thoroughly withMilli-Qwater. This wasfollowed by a similar electrochemical pre-treatment methoddescribed by Nagaoka and Yoshino [31]. Based on theRandles-Sevcik equation (Ipa¼ (2.69�105) n3/2AD1/2v1/2Cred),the electroactive area (A) of all electrodes was determinedfrom the anodic peak current (Ipa) obtained in the cyclicvoltammogram of 5.0 mM bulk concentration of Fe(CN)6

4�

(Cred) in 1.0 M KNO3 as supporting electrolyte at a specificpotential scan rate (v). Avalue of 6.5 � 10�6 cm2 s�1 at 20 8Cderived under similar experimental conditions by Kakuichiet al. [32] was used as the diffusion coefficient (D) ofFe(CN)6

4� and the number of electrons transferred, n, wasassigned as unity.To facilitate the binding of capture antibody, an electro-

chemically pretreated glassy carbon electrode was initiallyimmersed in a solutionof 1.5%(w/v)EDC in acetonitrile for2 h. This step was to activate the carboxylic groups on theelectrode by forming an o-acylurea intermediate [28]. Afterrinsing twice with acetonitrile the electrode was dried in anitrogen atmosphere. Prior to immobilising the captureantibody Ab1 on the electrode surface, the purchased stocksolution of Ab1 was washed twice with Tris buffer in aMicrocon filter device and then diluted to the concentrationof 750 mg mL�1 in Tris buffer, pH 7.6. A 10-mL aliquotcontaining 7.5 mg ofAb1was placed on the surface of the dryelectrode. The electrode was incubated at 4 8Covernight forthe reaction to proceed. Following Ab1 immobilisation,unbound antibody was removed by washing the electrodefive timeswith 100-mLaliquots of PBS. The antibody-coatedelectrodes were immersed in 5% (v/v) ethanolamine in PBSfor 10 min to block any unreacted but still active o-acylureaintermediates. After rinsing with water the electrodes wereready for use.

2.6. Determination of hCG by Sandwich ElectrochemicalImmunoassay

A sandwich electrochemical immunoassay was employedfor the determination of hCG in buffer and serum samples.A range of hCG doses in buffer or human serum wasincubated with the capture antibody, Ab1, immobilized onprepared electrodes as described above, for 30 min at roomtemperature. The electrodes were then washed thoroughlywith PBS and incubated with excess ferrocenecarboxylicacid-conjugated signal antibody (Ab2) in HEPES bufferfor15 – 20 min. The electrodes were rinsed with PBS beforeeach electrode was introduced in the electrochemical cell.Cyclic voltammetry in PBS was carried out between 0.1 andþ 0.6 V versus Ag jAgCl reference electrode. The concen-tration of hCG in serum samples from hospital patients wasdetermined using this sandwich immunoassay and quanti-fied by reference to an hCG calibration plot constructedusing hCG-free serum as the matrix.

2.7. Statistical Analysis

The statistical significance of all correlation coefficients andthe difference in hCG concentrations between our resultsand hospital results was evaluated at a 5% level usingStudentKs t test and the Wilcoxon signed-rank test, respec-tively.

239Signal Generation at an Electrochemical Immunosensor



3. Results and Discussions

3.1. Molecular Architecture of the Immunosensor

This work has been directed to developing an electro-chemical immunosensor using ferrocenecarboxylic acid asan electroactive label so that quantitative determination ofanalyte can be achieved by direct oxidation of this label. Thegeneral molecular architecture of our immunosensor isdepicted in Figure 1. To construct this immunosensor, Ab1was immobilized on a glassy carbon electrode to act as thecapture antibody for binding to its specific binding site(epitope) on our model analyte, human chorionic gonado-trophin (hCG). Following the addition of ferrocenecarbox-ylic acid-conjugatedAb2 that recognizes a different epitopeon hCG, it bound to the hCG already on the immunosensor.This sandwich format was employed to quantitativelydetermine the concentration of hCG in solutions.In this work, cyclic voltammetry for the oxidation of

ferrocenecarboxylic acid was performed so that the magni-tude of the resultant anodic peak current could be related tothe amount of hCG. Parameters involved in the preparationof an immunosensor will exhibit a direct influence on thevoltammetric response of the immunosensor. For example,the capacity for capturing hCGbyAb1will be dependent onthe quantity of Ab1 immobilized on the electrode surface.Therefore the immunosensor was optimized for amaximumvoltammetric response by determining the amount of Ab1immobilized on the immunosensor surface, incubation timein hCG solutions, and the amount of conjugate required forsandwich complex formation. Investigation of the effects ofthe latter two parameters was conducted when the electro-chemical immunosensor was used for the detection of hCGin human serum samples. The details in assembling each ofthe components and the application of the system foranalyte determination are described below.

3.2. A Glassy Carbon Electrode Based Immunosensor

In the development of electrochemical immunosensors, therecognition component of an immunocomplex is initiallyimmobilized on an electrode. A popular methodologyinvolves the formation of a self-assembled monolayer ofalkanethiols, such as 1-hexadecanethiol, 16-mercaptohex-adecanoic acid and thioctic acid, on a gold electrode [33 –

39]. The carboxylic acid groups of the monolayer are thencoupled to the amine groups of an antibody to form amidebonds via a carbodiimide reaction. Other components of animmunocomplex can then be securely anchored on theimmunosensor surface.On the other hand, the surface of glassy carbon electrodes

is well known to exhibit many reactive sites consisting ofsuch groups as hydroxylic, phenolic and carboxylic func-tionalities. Therefore, many strategies have been developedfor the chemical modification of the glassy carbon electro-des to provide a tailor-made surface with desirable charac-teristics for the analytical detection of targeted analytes[14,18,40,41]. In order to expose these functional groups,glassy carbon electrodes are often electrochemically pre-treated. For example, following pretreatment with a seriesof anodic and cathodic potential pulses, phenolic carboxylicand carbonyl groups on the electrode surface were identi-fied by X-ray photoelectron spectroscopy [42]. Such car-boxylic groups are appropriate direct coupling sites forantibody protein molecules. Another approach reported byLu et al. [18] involved the direct coupling of bacterialproteins, Protein A and Protein G, to the carboxylic groupson glassy carbon electrodes followed by antibody attach-ment to these bacterial proteins. Enhanced antibody bind-ing capacity was achieved at these immunosensors.In our work, we have initially adopted a pretreatment

method for glassy carbon electrodes as described byNagaoka and Yoshino [31], which involves a sequentialapplication of an anodic and a cathodic pulse. This is aimedat maximizing the functional groups on the electrodesurface, including the carboxylic acid groups, for bindingto a capture antibody during the first step in immunosensorconstruction. In the presence of pH 7.6 PBS, severalresearchers have reported the application of 1.8 V and�1.5 V (versusAg jAgCl reference electrode) as the anodicand cathodic potentials, respectively [18,31,42]. We havefurther examined thedurationof eachpotential pulse so thatmaximum functional groups could be obtained on theelectrode surface. In these experiments we have initiallykept the anodic potential at 1.8 V for 3.5 min and thecathodic potential of �1.5 V was applied over a periodbetween 0.5 – 2.5 min. Next, the anodic potential of 1.8 Vwas varied between 3.5 – 7.5 min, while the cathodic poten-tial was kept constant at�1.5 V for 0.5 min. Following eachapplication step, cyclic voltammetry at the electrode in PBSwas conducted between �0.4 V and 0.4 V. Representativebackground voltammograms obtained in these experimentsare shown in Figure 2. In general, the background currentswere observed to increase with prolonged anodization andcathodization duration. However, extended duration of theanodic potential resulted in amore pronounced effect on themagnitude of the background currents. According toNagaoka and Yoshino [31], oxygen functional groups,particularly phenolic oxygen and carbonyl oxygen, areincorporated on the pretreated glassy carbon surface duringthe anodization step. These changeswould then be expectedto yield a larger charging current between the potentiallimits used in the cyclic voltammetric experiments. In our

Fig. 1. A schematic representing the molecular architecture ofan immunosensor on a pretreated glassy carbon electrode.

240 M. Akram et al.



work, maximum background current was observed after a6.0-min anodisation and a 1.0-min cathodization. Furtherincrease of anodization and cathodization duration did notproduce any appreciably larger background current. Na-gaoka and Yoshino pointed out that an increase in back-ground current was a function of the number of carboxylicacid groups generated on the surface as a result of pretreat-ment of the electrode. Further, Engstrom et al. [42] dem-onstrated that there was no change in the electrode surfacearea resulting from the pretreatment, but an increase in theactivity of the functional groups that was responsible for theobserved increase in background currents. Thus anodizationat 1.8 V for 6 min and cathodization at �1.5 V for 1.0 minwere adopted for pretreating the glassy carbon electrodeused in subsequent experiments.

3.3. Activating Carboxylic Acid Groups on Glassy CarbonElectrode

In the next step of our immunosensor construction weemployed a carbodiimide reaction to attach an antibodyonto the immunosensor. This involves incubation of pre-treated glassy carbon electrodes in an EDC solution toactivate carboxylic acid groups. This has usually beenconducted by incubating an electrode in 1% (w/v) EDC(in anhydrous acetonitrile) for 5 h [28]. In our work we haveexamined the feasibility of using a shorter incubation timeby studying incubation periods varying from 30 – 300 min inboth 1% and 1.5% (w/v) EDC. This was determined bycyclic voltammetry of ferrocenecarboxylic acid conjugatedantibody immobilized on the EDC activated electrodes in

PBS. The results obtained are depicted in Figure 3.We showtwo cyclic voltammograms for the oxidation of 5.0 mMferrocenecarboxylic acid (in 0.1 M PBS as supportingelectrolyte) at a bare glassy carbon electrode and that forthe oxidation of ferrocenecarboxylic acid conjugated tosignal antibody on an activated electrode, respectively. Anoxidation peak at approximately 0.34 Vin both experimentsindicated that the signal obtained in the latter most likelycorresponded to the oxidation of ferrocenecarboxylic acidconjugated to Ab2. We then conducted similar experimentsafter incubating electrodes in 1.0% and 1.5%EDC solutionfor a range of duration. As shown in Figure 3(ii), there wasan expected higher current resulting from the oxidation offerrocenecarboxylic acid after the glassy carbon electrodewas immersed in 1.5%. EDC solution compared to that in1.0% solution. In both solutions, there was generally anincrease in the oxidation current as a function of immersionduration when this was less than 120 min with no significanteffect at longer times. Based on these results, the glassycarbon electrodes in all subsequent experiments wereactivated in 1.5% (w/v) EDC for 120 min prior to antibodycoupling.

3.4. Immobilization of the Capture Antibody

To quantify the amount of the capture antibody, we rely onthe determination of radioactivity from 125I-labelled Ab1immobilized on an electrode surface. In these experiments,an activated glassy carbon electrode was initially incubatedin a 125I-labelled Ab1 solution. Following adequate rinsing,the radioactivity of immobilized 125I-labelled Ab1 wasmeasured. Table 1 shows the binding efficiency of Ab1applied on the electrode surface. In general, enhancedbinding of Ab1 to the activated glassy carbon electrode wasobserved as the amount of Ab1 was increased from 0.5 to4 mg. The binding efficiency then leveled off at approx-imately 300 ng when more than 7.5 mg of Ab1 was applied.The relatively lowbinding ofAb1 to the surface is attributedto the fact that pretreating the electrodes not only exposedcarboxylic acid on the surface, but also graphitic oxides and�OHgroups formed as a result of the oxidation at the glassycarbon electrode surface [43]. However immunoglobulin G(Ab1) only chemically binds with the intermediate of anEDC reaction, o-acylurea, which is formed by reacting with

Fig. 2. Representative cyclic voltammograms obtained in PBS(pH 7.6) at a glassy carbon electrode following an anodization anda cathodization pretreatment at 1.8 Vand �1.5 V, respectively, forvarious times: constant 1.8 V for 3.5 min while �1.5 V was appliedfor a) 0.5 min and b) 2.5 min; constant �1.5 V for 0.5 min whileþ1.8 V was applied for c) 3.5 min and d) 7.5 min. Scan rate¼500 mV s�1.

Table 1. Amount of Ab1 bound following the application of anincreasing load of the capture antibody, Ab1, on a glassy carbonelectrode.

Amount of Ab1 applied (mg) Amount of Ab1 bound (ng) [a]

0.5 9� 11.0 22� 24.0 96� 87.5 300� 148.0 304� 148.5 298� 17

[a] All errors here represent standard deviations based on N¼ 3.




carboxylic groups on the surface. Other functional groupson the electrode surface would remain unreacted with EDCand were not capable of binding to Ab1. Notably, a large

amount of Ab1 on the pretreated electrode is generallyexpected to enhance the electrochemical activity achievableat an immunosensor, which will aid in improving the

Fig. 3. (i) Cyclic voltammograms for the oxidation of a) 5 mM ferrocencarboxylic acid (in 0.01 mol L�1 PBS as supporting electrolyte)at a bare glassy carbon electrode, b) ferrocenecarboxylic acid conjugated to the signal antibody Ab2 on an immunosensor. Scan rate¼50 mV s�1; (ii) Electrochemically pretreated glassy carbon electrodes were immersed in 1.0% (^) and 1.5% (w/v) (&) EDC solution,respectively, for various time periods shown. The current density obtained for the oxidation of ferrocenecarboxylic acid immobilized onthe electrode was then plotted against the immersion duration.

242 M. Akram et al.



detection limit of hCG. Based on the above, 7.5 mg of Ab1were employed in our work.

3.5. The Model Analyte, Human ChorionicGonadotrophin

Our immunosensor was designed to measure the pregnancyhormone and tumor marker hCG. This is a 37-kDaglycoprotein hormone produced by the normal trophoblastsof the placenta during pregnancy and by trophoblastictumor cells [44, 45]. Shortly after conception, hCG isproduced by the developing placenta and increases through-out the first trimester, doubling every one to two days. Alevel of 25 IU L�1 in serum or urine is usually regarded as anindicator of early pregnancy [46] (1 mg hCG¼ 9825 Inter-national Units (IU) of hCG). Trophoblastic tumors thatsecrete hCG include gestational and testicular choriocarci-noma where levels of hCG in serum or urine may reachhundreds to thousands of International Units per litre. ThehCG levels can be used to determine tumormass, successfultreatment of malignancy or recurrence of disease. Structur-ally, each hCGmolecule consists of two dissimilar subunits,termed a and b [47]. The a-subunit is identical to that of thepituitary hormones, luteinizing hormone, follicle stimulat-ing hormone, and thyroid stimulating hormone, while the b-subunit is specific to hCG. Antibodies that are directedtoward theb-subunit enable the specific detection of hCG tobe made by immunoassay. In the construction of thisimmunosensor, specificity for hCG was achieved by usingtwo monoclonal antibodies that were directed towarddifferent binding sites on the hCG molecule.

3.6. Immunosensor Performance

Having completed the study of each component in themolecular architecture, we assessed the ability of theimmunosensor to quantify hCG in buffer solutions and inhuman serumby cyclic voltammetry at the immunosensor in10 mM PBS. Figure 4 shows the voltammetric responseobtained with hCG prepared in 0.5% BSA-PBS over aconcentration range up to 3,200 IU L�1. In each case, whenthe potential was scanned between 0.0 Vand þ 0.6 V, awell-defined voltammogram consisting of an oxidation peakbetween 0.34 – 0.37 V, and a reduction peak at 0.25 – 0.28 Vin the reverse scan was obtained. Unlike enzyme labels, theoxidation current was obtained almost instantaneously inour experiments with no delay needed for equilibration andincubation time. Further, no stringent temperature controlwas required. More significantly, direct oxidation of ferro-cenecarboxylic acid conjugated to Ab2 has yielded thevoltammetric response. A linear calibration plot (alsoshown in Fig. 4) was then constructed based on oxidationpeak current density in order to normalize variationsbetween electrodes. The plot in Figure 4(iii) is representedby the equation, current density/mA cm�2¼ [hCG]/IU L�1

0.1186 (�0.0059)þ 0.0005 (�0.0006) (where the errors here

Fig. 4. (i) – (ii) Cyclic voltammetric responses of immunosensorsobtained in 0.5% BSA-PBS spiked with different concentrationsof hCG solutions. Scan rate¼ 50 mV s�1; (iii) calibration graphobtained by plotting the oxidation current density in (i) and (ii)against the concentration of hCG.




refer to the 95% confidence intervals) (correlation coef-ficient¼ 0.9985 and this was found to be statisticallysignificant at a 95% level based on N¼ 9). These resultsindicated the feasibility in relying on the direct oxidation offerrocenecarboxylic acid for quantification of hCG at ourimmunosensor.

3.7. Matrix Matched Dose-Response Plot for theDetermination of hCG

Serum comprises a complex matrix of various proteins,lipids and sugars. Therefore, the performance of ourimmunosensor in the determination of hCG was alsoevaluated in serum samples. In these experiments, aliquotsof human serumcontaining nohCGwere spiked to give finalhCG concentrations between 50 and 3200 IU L�1. Then,10 mL of each of the spiked solutions were placed on animmunosensor with Ab1 already immobilized on theelectrode. After 30 min each immunosensor was rinsedwith PBS to remove unbound hCG. Next, 8.5 mL conjugatecontaining 2.5 mg Ab2 were placed on the surface of theimmunosensor for 15 to 20 min at room temperature. Afterthorough rinsing with PBS to remove excess unboundconjugate, each immunosensor was placed in an electro-chemical cell containing pH 7.6 PBS. Cyclic voltammetrywas carried out between 0.0 Vand 0.6 V. Figure 5 shows thevoltammetric results obtained when various hCG concen-trations were spiked in blank human serum as calibrationstandards, as well as the linear calibration plot obtained.This calibration plot is represented by the equation, currentdensity/mAcm�2¼ 0.1000 (�0.0046) [hCG]/IUL�1þ 0.0323(�0.0015) (correlation coefficient¼ 0.9987 and this wasfound to be statistically significant based on N¼ 8), with adynamic range up to 1,500 IU L�1 was obtained. Based onthree and ten times the standard deviation of a blank value,the limit of detection and the limit of quantification wereestimated to be 2.2 and 2.5 IU L�1 hCG, respectively. When25 mL of conjugate containing 7.5 mg of Ab2 were immo-bilized on the immunosensor surface, we were able toextend the dynamic range of the analyte concentration to3,200 IU L�1. However when additional quantities ofconjugate were applied the dynamic range remainedunchanged. We attribute this result to a limited ability ofour immobilization strategy to orientate the immunogenicdomain of the capture antibody towards the targeted hCGanalyte. In turn, this led to no further binding on addition oflarger quantities of Ab2, the signal antibody.Next, a comparison of hCG levels in serum samples using

our immunosensor was conducted with those reported by ahospital diagnostic laboratory and this is shown in Table 2.Based on the Wilcoxon signed ranked test, no significantdifference was found between the group of experimentalresults and the group of hospital values at a 5% level.However, due to the high dose hook effect, serum sampleswith a high hCG level must be diluted before analysis usingour immunosensor. For example, a 100 times dilution wasnecessary in analysing Sample 7 in Table 2 to obtain a signal

Fig. 5. (i) – (ii) Cyclic voltammetric responses of immunosensorsobtained in hCG free human serum spiked with differentconcentrations of hCG solutions. Scan rate ¼ 50 mV s�1; (iii)calibration graph obtained by plotting the oxidation currentdensity in (i) and (ii) against the concentration of hCG.

244 M. Akram et al.



within the current calibration range. Owing to the amplifiederrors involved in multiple dilution steps, there was asignificant difference between our value and the hospitalvalue for Sample 7. As a result, at this stage of development,we regard our immunosensor to be useful for the detectionof low levels of hCG such as those found in early pregnancy.

4. Conclusions

In this work, ferrocenecarboxylic acid was applied as anelectroactive label on an electrochemical immunosensor.The oxidation of this label was then demonstrated to befeasible in the development of a direct signal-generatingscheme at the electrochemical immunosensor. In contrast tocommonly employed schemes involving an enzyme label,there was no equilibration and incubation time required forthe oxidation of ferrocenecarboxylic acid. Moreover, thereaction was simply carried out at room temperature inreadily available buffer solutions. When the immunosensorwas applied to the detection of hCG as a model analyte, alinear calibration range up to 3,200 IU L�1 hCG and adetection limit (based on three times the standard deviationof the blank) of 2.2 IU L�1 of hCG were achieved. Theseresults show that our immunosensor will be useful for thedetermination of hCG levels in early pregnancy. In addition,statistically acceptable agreements were achieved in thehCG levels present in several serum samples analyzed byour laboratory compared to those reported by a pathologylaboratory in a Sydney hospital.

5. References

[1] P. B. Luppa, L. J. Sokoll, D. W. Chan, Clin. Chim. Acta 2001,314, 1.

[2] O. A. Sadik, Electroanalysis 1999, 11, 839.[3] H. J. Cruz, C. C. Rosa, A. G. Oliva, Parasitol. Res. 2002, 88,

s4.[4] D. A. Palmer, J. N. Miller, Anal. Chim. Acta 1995, 303, 223.[5] C. A. Wijayawardhana, H. B. Halsall, W. R. Heineman, In

Electroanalytical Methods for Biological Materials (Eds: A.Brajter-Toth, J. Q. Chambers), Marcel Dekker, New York2002, pp. 329 – 365.

[6] C. A. Wijayawardhana, H. B. Halsall, W. R. Heineman, InEncyclopedia of Electrochemistry, Vol. 9 (Ed: G. S. Wilson),Wiley-VCH, Weinheim 2002, pp. 148 – 150.

[7] A. Kokado, A. Tsuji, M. Maeda, Anal. Chim. Acta 1997, 337,335.

[8] Z.-Z. Li, F.-C. Gong, Sens. Actuators 2004, B 99, 562.[9] N.-Y. Pan, J.-S. Shih, Sens. Actuators 2004, 98, 180.

[10] S. Lçfas, B. Johnsson, J. Chem. Soc., Chem. Commun. 1990,21, 1526.

[11] B.-K. Oh, Y.-K. Kim, W. Lee, Y. M. Bae, W. H. Lee, J.-W.Choi, Biosens. Bioelectron. 2003, 18, 605.

[12] T. C. Tang, A. P. Deng, J. H. Huang, Anal.Chem. 2002, 74,2617.

[13] T.-K. Lim, N. Nakamura, T. Matsunaga, Anal. Chim. Acta1998, 370, 207.

[14] Y.-T. Kong, M. Boopathi, Y.-B. Shim, Biosens. Bioelectron2003, 19, 227.

[15] S.-Q. Liu, H.-X. Ju Anal. Biochem. 2002, 307, 110.[16] M. Santandreu, S. Alegret, E. Fabregas, Anal. Chim. Acta

1999, 396, 181.[17] M. P. Kreuzer, M. Pravda, C. K. OKSullivan, G. G. Guilbault,

Toxicon 2002, 40, 1267.[18] B. Lu, M. R. Symth, J. Quinn, D. Bogan, R. O. Kennedy,

Electroanalysis 1996, 8, 619.[19] B. Law, Immunoassay A Practical Guide, Taylor & Francis,

London 1996, 126.[20] F. J. Hayes, B. H. Halsall, R. W. Heineman, Anal.Chem. 1994,

66, 1860R.[21] M. Yasuzawa, T. Matsuki, H. Sako, A. Kungi, Chemical

Sensors 1998, 14 (suppl. B), 65.[22] T.-K. Lim, T. Matsunaga, Biosens. Bioelectron. 2001, 16, 1063.[23] M. Okochi, H. Ohta, T. Tanaka, T. Matsunaga, Biotech.

Bioeng. 2005, 90, 14.[24] R. Ojani, J. B. Raoof, A. Alinezhad, Electroanalysis 2002, 14,

1197.[25] H. S. Mandal, H.-B. Kraatz, J. Organomet. Chem. 2003, 674,

32.[26] M. Nakayama, T. Ihara, K. Nakano, M. Maeda, Talanta 2002,

56, 857.[27] K. D. Gleria, H. A.O. Hill, C. J. Mcneil, Anal. Chem. 1986,

58, 1203.[28] C. Duan, M. E. Meyerhoff, Mikrochim. Acta 1995, 117, 195.[29] F. C. Greenwood, W. M. Hunter, Biochem. J 1963, 89, 114.[30] C. A. Brutis, E. R.Ashwood, Tietz Textbook of Clinical

Chemistry, W. B. Saunders, PA, USA 1994, 834.[31] T. Nagaoka, T. Yoshino, Anal. Chem. 1986, 58, 1037.[32] T. Kakiuchi, H. Usui, D. Hobara, M. Yamamoto, Langmuir

2002, 18, 5231.[33] Y. Dong, C. Shannon, Anal. Chem. 2000, 72, 2371.

Table 2. hCG levels determined by our immunosensor against the values obtained by a Sydney hospital.

Sample Experimental values [a] IU L�1 Hospital values IU L�1 Percentage difference

1 109� 1.0 112 �2.6%2 47� 1.5 55 �14.5%3 12.3� 2.0 10.7 þ 14.9%4 150� 4.4 148.6 þ 0.9%5 386� 23.8 366.6 þ 5.3%6 4.7� 1.5 4.6 þ 2.1%7 45000� 3560 67000 �32.8%8 <2.2 Not detectable –

[a] All errors here represent standard deviations based on N¼ 3




[34] M. Dijksma, B. A. Boukamp, B. Kamp, W. P. Bennekom,Langmuir 2002, 18, 3105.

[35] D. M. Disley; D. C. Cullen; H.-X. You, C. R. Lowe, Biosens.Bioelectron. 1998, 13, 1213.

[36] J. M. Clyne, J. A. Running, M. Stempien; R. S. Stephens, H.Akhavan-Tafti, A. P. Schaap, M. S. Urdea J. Biolumin.Chemilumin. 1989, 4, 357.

[37] T. M. OKRegan, L. J. OKRiordan, M. Pravda, C. K. OKSulli-van, G. G. Guilbault, Anal. Chim. Acta 2002, 460, 141.

[38] J. Rickert, W. Gopel, W. Beck, G. Jung, P. Heiduschka,Biosens. Bioelectron. 1996, 11, 757.

[39] J. C. OKBrien, V. W. Jones, M. D. Porter, Anal. Chem. 2000,72, 703C.

[40] C. Peihong, R. L. McCreery, Anal. Chem. 1996, 68, 3958.[41] C. Saby, B. Ortiz, Y. G. Champagne, D. Belanger, Langmuir

1997, 13, 6805.[42] R. C. Engstrom, V. A. Strasser, Anal. Chem. 1984, 56, 136.[43] A.Dekanski, J. Stevanovic, R. Stevanovic, B. Z. Nikolic,

V. M. Jovanovic, Carbon 2001, 39, 1195.[44] G. D. Braunstein, J. Rasor, D. Alder, H. Danzer, M. Wade,

Am. J. Obstet. Gynecol. 1976, 126, 678.[45] G. B. Pastofide, D. P. Goldstein, T. S. Kosasa, Am.J. Obstet.

Gynecol. 1974, 120, 1025.[46] L. A. Cole, J. Reprod. Med 1998, 43, 3.[47] S. J.Segal, Chorionic Gonadotrophin, Plenum Press, New

York 1980, 507 – 520.

246 M. Akram et al.



Documents

Signal Generation at an Electrochemical Immunosensor via the Direct Oxidation of an Electroactive Label