This article appeared in a journal published by Elsevier. The attached copy is furnished to the author for internal non-commercial research and education use, including for instruction at the authors institution and sharing with colleagues. Other uses, including reproduction and distribution, or selling or licensing copies, or posting to personal, institutional or third party websites are prohibited. In most cases authors are permitted to post their version of the article (e.g. in Word or Tex form) to their personal website or institutional repository. Authors requiring further information regarding Elsevier’s archiving and manuscript policies are encouraged to visit: http://www.elsevier.com/copyright
Transcript
This article appeared in a journal published by Elsevier. The attachedcopy is furnished to the author for internal non-commercial researchand education use, including for instruction at the authors institution
and sharing with colleagues.
Other uses, including reproduction and distribution, or selling orlicensing copies, or posting to personal, institutional or third party
websites are prohibited.
In most cases authors are permitted to post their version of thearticle (e.g. in Word or Tex form) to their personal website orinstitutional repository. Authors requiring further information
regarding Elsevier’s archiving and manuscript policies areencouraged to visit:
a Institute of Microbial Technology (CSIR), Sector 39-A, Chandigarh 160036, Indiab Swiss Federal Institute of Technology, Lausanne EPFL IC ISIM LSI1, INF 333 (Bâtiment INF), Station 14, CH-1015 Lausanne, Switzerland
a r t i c l e i n f o
Article history:Received 2 September 2011Received in revised form 4 November 2011Accepted 14 November 2011Available online 25 November 2011
It has been already proved that electrical capacitance increases at the electrode surface due to bind-ing of highly charged antigens on the biointerface. In this paper, this capacitance change was used todemonstrate a novel detection for cardiac troponin I (cTnI). Citrate-capped gold nanoparticles (GNPs)were deposited on screen printed electrodes using an innovative one-step electrochemical technique.The gold (Au) nanostructuration onto the electrode surface was characterized by atomic force and fieldemission scanning electron microscopies. The highly specific anti-cTnI antibodies were immobilized ontothe Au matrix, and characterized by confocal microscopy. The interaction of cTnI with its correspondingantibody was studied with respect to electrical capacitance changes as deduced by using the Randlesmodel of the equivalent circuit of the electrochemical cell. The sensor-to-sensor measurements errorwas registered less than 2% on the average measured capacitance. cTnI was successfully detected in alabel-free manner with a limit of detection equal to 0.2 ng mL−1. This obtained result was one order ofmagnitude better than that obtained with ELISA tests performed by using the same antibodies, with adetection limit of 4.3 ng mL−1.
The design of label-free, fast, reliable, cost effective, affinity-based diagnosis is one of the objectives in current biochemicalresearch. The trend in electrochemical immunosensors is fastshifting from enzyme-linked detection of redox active compoundsto direct detection of changes in physical parameters, such asimpedance, capacitance, resistance, optical properties or masschanging during antigen–antibody interactions. The idea is todevelop smart microprocessor-controlled portable-devices forpoint-of-care testing of clinically relevant disease biomarkers[1–3]. Cardiac markers represent the leakage of cell contents fromcardiac myocytes. This release of cardiac markers happens whenthe heart muscles are damaged. An elevated level of those markersusually indicates myocardial infarction. There are different kinds ofbiomarkers for cardiac injury, such as creatine-kinase, lactate dehy-drogenase isoenzymes, myoglobin, troponin T & I, but only few areroutinely used in hospital practice [4]. Troponin I is one of those.Human cardiac troponin I is one of the three subunits of the cardiactroponin complex. It binds to actin in thin myofilaments to supportthe troponin–tropomyosin complex and it inhibits contraction. cTnI
is usually only contained inside the heart muscle and it is currentlyconsidered to be the standard biomarker test for detecting acutemyocardial infarction, because it is significantly more specific thanother heart markers [5]. cTnI levels begin to rise 2–3 h after themyocardial infarction and elevation in levels can persist for up to10 days, making it ideal for retrospective diagnosis of infractions[6,7].
Various immunodetection methods are currently used fortroponin I detection [8,9]. However, these methods are time-consuming and they usually require labeled reagents andbulky instrumentation. The label-free immunodetection based onimpedance spectroscopy is currently gaining importance becauseof their simplicity and fast response time. For these reasons, it isespecially suitable for developing point-of-care devices [10–12].The most common transducer platforms are based on electrodeswith interdigitated finger geometries. Self-assembled monolayers(SAMs) including ethylene glycol alkanethiols [13] or Protein A [14]or Protein G [15], are used to anchor the antibody on the sensorsurface in order to stabilize the biointerface. However a major con-cern is about the costly nature of the used linker molecules becausethe costs of those chemicals and process to get self-assembled pre-cursor or blocking films. An alternative to decrease the costs is touse GNPs, that improve the performances of electrodes [16]. GNPsare suitable because they are produced with easy methodologies, incontrolled monodisperse suspensions. Moreover, they present veryhigh surface-to-volume ratio that results in a (more loading) when
762 V. Bhalla et al. / Sensors and Actuators B 161 (2012) 761– 768
dealing with antibody binding. They also provide a field enhance-ment at the surface interface [17,18]. Several reports have showngold nanostructuration in applications to biosensing. Drop-castingis a simple way of obtaining a GNP modified electrodes [19]; but itis mainly possible for nanoparticles in volatile organic solvents andnot in case of water disperse colloidal solution. Direct GNPs elec-trodeposition has been carried out for impedance based monitoringof antigen–antibody interactions [20]. However, any direct reduc-tion of tetrachloroauric acid (HAuCl4) on electrodes would justresult in bulk deposition of gold. Nanostructuration with citrate-capped gold nanoparticles has been previously carried out witha different approach by using electrostatic assembly of negativelycharged particles onto positively charged amine modified electrodesurfaces [21,22]. Although it leads to immobilization of GNPs, thelinker layer used for the generation of the amine functional groupsmainly provides an insulating layer at the electrode surface.
In the present manuscript we use GNPs for creating a nanosizedcolloidal Au base matrix onto screen printed electrodes (SPE) inorder to propose a low-cost, one-step, electrochemically controlledimmobilization platform for diagnosis purposes. The innovation inthe proposed method relies on the dual role of the colloidal goldGNPs as a matrix for direct antibody immobilization and transduc-tion properties. They lead to applications in sensitive detection ofcTnI in a label-free capacitive manner by using the highly chargednature of the antigens at physiological pH.
2. Materials and methods
2.1. Reagents
Troponin I monoclonal antibody (mAb), recombinant humancardiac Troponin I antigen were obtained from Biospacific, USA.Gold chloride salt, citric acid, rhodamine isothiocyanate (RITC) andalkaline phosphatase conjugated antibody (anti-cTnI polyclonal)were obtained from Sigma–Aldrich. Para-nitrophenyl phosphate(pNPP) substrate for alkaline phosphatase was procured from Ban-galore Genei, Bangalore, India. Maxisorb polystyrene Nunc plateswere obtained from Fisher Scientific and used for conventionalELISA tests. All the other reagents were acquired from Qualigens,India. Water used was purified with a Milli-Q ultra pure sys-tem (Millipore, India) and its final resistivity was smaller than18 M� cm−1.
2.2. Preparation of colloidal gold base matrix immunosensors
The monodispersed colloidal gold nanoparticles were preparedby slight modification of Frens method [23]. 100 mL of 0.01% solu-tion of HAuCl4 in deionised water was taken in 250 mL coveredErlenmeyer flask. The solution was brought to boiling point on ahot plate stirrer and 4 mL of 1% sodium citrate solution was rapidlyadded to the boiling solution. The color of the solution changed tobright wine red, an indicative of the formation of gold nanoparti-cles. The solution was allowed to boil for another 10 min and, then,cooled down for storing at 4 ◦C. The final average size of the particleswas characterized using UV–Vis spectroscopy (Analytikjena S600)and HR-TEM (FEI-TECNAI G2 F20) and it was found to be equal to20 ± 5 nm.
The citrate-capped GNPs were diluted 1:5 times in the 10 mMphosphate buffer, pH 7.4 and 100 �L droplet was spread onto theSPE surface completely covering the three electrodes. The elec-trodeposition of the negatively charged citrate-capped GNPs wasperformed by using bulk electrolysis technique with an appliedpositive potential of 0.7 V vs. Ag/AgCl for 1500 s. The immobiliza-tion of antibodies was carried out overnight, incubating 25 �g mL−1
of monoclonal antibodies in 10 mM phosphate buffer (pH 7.4), at
4 ◦C. The electrode surface loaded with antibody molecules wasrinsed with phosphate buffer (pH 7.4) to remove the weakly boundimmunoglobulins.
2.3. Sensor characterization by AFM, SEM and confocalmicroscopy
Atomic force microscopy (AFM) of the electrodes was carriedout using solver pro 7 model (NTMDT, Russia). All the studies wereperformed in non-contact mode using NSG01 “Golden” silicon can-tilevers (NTMDT) having a typical spring constant of 5.1 N m−1.The cantilever where drive at resonance frequencies were from150 up to 160 Hz. The images were acquired with a scan rate of1.01 Hz and background was subtracted by using the Nova imageanalysis software (NTMDT). The AFM imaging was done on flatgold surfaces evaporated onto silicon substrates in order to suc-ceed in AFM investigations otherwise impossible on SPE surfacesdue to their roughness at the �m scale. Imaging on SPEs wereindeed directly performed with field emission scanning electronmicroscope (FESEM). For that, the samples were mounted on analuminum stub by using double adhesive carbon tape. The FESEMinstrument was Hitachi S-4300 and it was used at an acceleratingvoltage of 10 kV.
Confocal microscopy was used to characterize the biointerfacedeveloped for the immunosensor aim by labeling the antibodymolecules with rhodamine isothiocyanate (RITC) fluorophore. Theantibody solution (1 mg mL−1) was first dialyzed against carbonatebuffer (50 mM, pH 9.2) and stock solution of RITC (0.5 mg mL−1)prepared in carbonate buffer was used for preparing RITC labeledantibody conjugates (F:Ab). The fluorescent antibodies on thesensor surface were investigated by Carl Zeiss inverted confocalmicroscope, using 400× magnification.
2.4. Electrochemical impedance measurements
The impedance measurements were performed by using anelectrochemical analyzer (CH instruments 660-D) with multi-plexer CHI 684-8. The printed electrodes from Zensor (modelTE100) were acquired from CH instruments, USA. They consistof a 3-mm-diameter carbon-paste working-electrode, a counter-electrode made of carbon and an Ag/AgCl reference-electrode. Theelectrodes were mechanically polished with 1.0, 0.3 and 0.05 �m�-Al2O3 powder and successively washed ultrasonically in distilledwater. Electrochemical impedance spectroscopy data was acquiredon these electrodes over a frequency range from 1 Hz up to 100 kHz.The measures were repeated after formation of colloidal Au matrix,after immobilization of the antibodies over colloidal Au surface andafter incubation of TnI antigen over the electrode. The amplitudeof the alternating voltage was 5 mV without any DC bias voltage.All the measurements were recorded in dark, at room tempera-ture (25 ◦C). All the analysis were carried in PBST buffer (phosphatebuffer saline solution, water solution 137 mM NaCl, 100 mM phos-phate, 2.7 mM KCl, at pH 7.4, supplemented with 0.05% tween-20 asdetergent). The data was analyzed using standard system softwareavailable with the service software in the CH workstation.
2.5. Sandwich ELISA
The Nunc microtiter plates were incubated with 5 �g mL−1 ofcapture antibody (anti-cTnI) overnight at 4 ◦C. After washing threetimes with PBS, the blocking treatment was carried out with 10%skim milk for 1 h at 37 ◦C. The plate was washed three timeswith PBS and with PBST (once) followed by incubation with cTnIantigen at various concentrations for 1 h at room temperature.Subsequently, 100 �L of alkaline phosphatase conjugated antibody(anti-cTnI polyclonal) at a dilution of 1:1000 was added to each of
V. Bhalla et al. / Sensors and Actuators B 161 (2012) 761– 768 763
Fig. 1. CV scans of SPE in PBS buffer, pH 7.4, before and after electrodeposition ofthe GNPs. The dotted line shows the CV scan of bare SPE and the solid line showsafter GNP modification. The scans were obtained at a scan rate of 50 mV s−1. Thereference electrode was screen printed Ag/AgCl. Inset: Bulk electrolysis at 0.7 V fordeposition of gold nanoparticles. The relative current is shown with respect to thedeposition time.
the wells and incubated at 37 ◦C for 1 h. After a single wash with100 �L of pNPP, the substrate for alkaline phosphatase was added toeach well and the plate was incubated for 15 min at 37 ◦C followedby OD (optical density) measurement at 405 nm with multimodereader (Synergy Biotech, USA).
3. Results and discussion
3.1. Morphological characterization of colloidal gold matrix andproper biointerface development
The synthesized Au colloid nanoparticles were characterized byUV–Vis spectroscopy and transmission electron microscopy (TEM)(data shown in Fig. S1 in supplementary data). A maximum absorp-tion in UV–Vis spectrum was observed at 520 nm, which is a typicalplasmon absorbance of spherical gold nanoparticles. The antibod-ies tagged with GNPs showed a slight shift (∼10 nm) confirmingthe formation of antibody–GNP complex. Transmission electronmicroscopy confirms that the particles size was approximately of20 nm. The relative current during the electrochemical depositionof GNP decreased with respect to time. This current decrease con-firms the deposition and it becomes stable after 1500 s (inset toFig. 1). The cyclic voltammetry (CV) scans for the deposited GNPsshow an enhanced surface area for the immobilization of anti-bodies (Fig. 1). The gold peaks are more resolved in 0.1 M acetatebuffer (pH 5.2), (scans are presented in supplementary information
by Figs. S2 and S3). The supplementary information also show CVscans obtained for another electrode type LC-LAGE (low-cost laserablated gold electrodes) described previously in our report [24].
The AFM study was carried out by depositing the GNPs on aflat gold surface, fabricated onto a silicon chip. In Fig. 2A, the AFMmicrographs show both the gold terraces and the spherical featuresrelated to our nanoparticles. The x–y cross section (Fig. 2B) clearlyindicates that the size of the individual features in the AFM image isclose to the expected nanoparticles size. The nanoparticles provideelectric field enhancement that completely overlap the biomolec-ular interaction at the nanoscale that assures a biosensing with asensitivity better than bulk gold [25,26]. The size dependence ofGNP electrical properties was discussed previously [27]. It is gen-erally agreed that an average size that is not too large (towardsbulk materials) and not too small (towards quantum effects) isa better compromise to get the various GNP benefits. The opti-mized average size has been verified ranging from 15 up to 30 nm,a size which is closer to the antibodies size. The gold modified sur-faces were analyzed for root mean square (RMS) roughness analysisby AFM on a representative area of 5 �m2. The RMS roughnessfor bare gold surface was 0.73 ± 0.1 nm. The RMS increases up to9.95 ± 1.1 nm for the GNP electrodeposited surface. This increaseindicates a dramatic increase of the surface area, up of one orderof magnitude. GNP assembled onto the surface of SPE can enhancethe amount of antibody molecules immobilized on the electrode,and preserve their immune activities as well [28]. The adsorptionof antibodies onto the GNP surface is established by a bondingbetween the –NH3
+ and COO− functional groups [29]. The bondformation is supported by the electrostatic interaction between thesurface-terminated anionic groups (–COO−) on the GNPs, and thepositively charged amine groups (–NH3
+) on the lysine residuesof the antibody. The Au–NH2 bond is another important bond thatleads to the immobilization of antibody on the GNP surface [30–33].Antibodies bound to the nanostructured electrodes are physicallyflexible and suitable for preventing the loss of protein conforma-tion or biological activity [34]. The RMS roughness further increaseafter the antibody immobilization on the GNP layer. The analysisof the SEM images confirm the features found with the AFM inves-tigations. Fig. 3 shows the FESEM images of colloidal Au modifiedscreen printed carbon electrodes obtained by electrodeposition atthe potential of 700 mV (vs. Ag/AgCl) for 1500 s. The FESEM imagesshown in Fig. 3A–C at different magnifications revealed colloidalAu. They are evident as bright spherical objects against a dark back-ground. The particles are randomly distributed on the surface of theelectrodes.
The efficient antibody immobilization on the GNP surface wasshown using confocal microscopy techniques by labeling the anti-body with rhodamine isothiocyanate (RITC) (Fig. 4A). The data showlight fluorescent regions and dark patches. The dark patches cor-respond to the surface of the modified electrodes left unoccupied
Fig. 2. AFM of the colloidal Au modified electrode. Square size is 250 nm2. (A) 3D image and (B) x–y cross-section of the image.
764 V. Bhalla et al. / Sensors and Actuators B 161 (2012) 761– 768
Fig. 3. FESEM image of the GNPs on the SPE surface at different magnifications. The white spots in the image correspond to the GNPs and the darker regions correspond tothe carbon electrode surface. The scale bar corresponds to 50 nm.
Fig. 4. (A) Confocal microscopy images of the RITC labeled antibodies immobilized on the nanosized Au colloid modified SPE surface. The surface contour of the modifiedelectrodes regions marked in the picture is also shown along the axis of the image. (B) Color-coded image: the red color represent immobilized antibodies and yellowrepresents the regions without immobilization of antibodies. (For interpretation of the references to color in this figure legend, the reader is referred to the web version ofthe article.)
by the antibodies. The figure also shows the contour (shown alongthe image axis) of the GNP-mAb modified screen printed electrodesurface. As can be clearly seen in the color-coded image (Fig. 4B),the surface is extensively modified with the antibodies, indicatinggood immobilization. In Z splicing mode, the fluorescent signals onthe SPE surface were recorded at a total 0–8 �m (supplementaryinformation Fig. S4).
3.2. Monitoring of immunochemical reactions
Impedance based methods provide information about the elec-trochemical interfacial properties of the modified surface of theelectrodes. The data reported in this paper were recorded by adding100 �L of PBS-T buffer, pH 7.4 on the SPE surface. No redox probesor labels were used in these experiments. The absence of a redoxprobes and, then, the subsequent non-faradic behavior is of advan-tage in using for blood serum where a multitude of other materialscan interfere with true faradic response. The data recording in PBSis also better since ferricyanide solution, used for faradic mea-surements, acts as mild oxidant that can denature some proteins[35,36]. Nyquist plots from the electrochemical impedance spec-troscopy after the various surface modifications and the modifiedRandle’s equivalent circuit [37,38] are shown in Fig. 5. The circuitincludes the following five elements: (i) the ohmic resistance ofthe electrolyte solution, Rsol; (ii) the Warburg impedance, ZW; (iii)capacitance of double layer (Cdl), associated with the double-layer,reflecting the interface between the assembled film and the elec-trolyte solution. In general, the latter element can also be replacedby the constant phase element (CPE) impedance, ZCPE [39]. The
circuit is completed by (iv) the electron transfer resistance, RCT [40]and (v) CPE [41]. The use of a CPE element in series, provided a goodfit to the data which is also logical, as we are not blocking ion path-ways to the electrodes, as in one of our previous reports, wherewe used ethylene glycol SAM to protect the electrodes from ions inthe solution [42]. The impedance data were fitted by using instru-ment manufacturer’s software, and the fitting parameters related
Fig. 5. Nyquist plot obtained after GNP deposition and after overnight incubationof the electrodes with anti-cTnI monoclonal antibodies. The plot after binding ofantigen to the antibodies is shown as well. The impedance spectra were taken at pH7.4 in PBS buffer and in the frequency range from 1 Hz to 100 kHz. Inset shows theRandles equivalent circuit used for fitting the data.
V. Bhalla et al. / Sensors and Actuators B 161 (2012) 761– 768 765
Table 1Major electrical parameters variation associated with biomolecular activity at the GNP modified surface. The scans were obtained at pH 7.4 in PBS-T buffer.
0.1 ng mL−1 225.5 518 8038 2.37 7.235 0.031250.39 ng mL−1 226.1 533.9 13 350 2.338 7.557 0.029420.78 ng mL−1 227.3 536.9 13 350 2.017 7.556 0.03341.56 ng mL−1 229.5 551.8 13 060 2.248 7.73 0.029013.125 ng mL−1 228.1 557 17 510 2.373 7.912 0.026366.25 ng mL−1 228.6 564.5 25 170 1.765 7.983 0.0333412.5 ng mL−1 227.8 569.8 23 990 2.185 8.181 0.02851
to the Randles’ model are shown in Table 1. The good fitting over theentire measurement frequency range indicates the good agreementbetween the circuit model and the measurements.
Fig. 5 mainly shows very different plots for the bare SPE surfaceand for the various surface modifications. The origin of the plots wasdetermined by the solution resistance (Rsol) [37,38]. The shape andseparation between plots are determined by the electron transferresistance Rct as well as the double layer capacitance (Cdl). The Rct
is determined by the modified electrodes conductivity while theCdl is related to double-layering phenomena [41]. The same ori-gin of the plots indicates no change in Rsol. Indeed, the Cdl and Rct
change, mainly increasing, as we immobilize the antibodies ontothe GNP-modified surface or when the antigen–antibody interac-tion takes place. Ideally, Rsol and ZW represent the properties of theelectrolyte solution and diffusion of the redox probe and, thus, arenot affected by modifications occurring on the electrode surface[43]. Inferences from the antigens are only possible on Rct and Cdlparameters. A negligible change in Rsol, ZW and CPE was observedin our experiments for the non-faradic impedance record during
the different surface modifications. At the same time, the changesin Cdl and Rct showed significant variations (see Table 1). The Rct
variation only depends on the larger occupancy of the antigen ontothe surface. On the other hands, the Cdl is a more suitable parameterfor sensing because it is affected by the antigen charge. The bind-ing of antigen (cTnI) to its monoclonal antibody leads to chargingof the GNP/mAb modified SPE surface and an increase in capaci-tance is observed. See Scheme 1 for an illustrative explanation ofthe phenomena. This increase in capacitance is consistent with ourprevious data on self assembled monolayers and interdigitated fin-ger geometry electrodes [44], as due to the double layer structure atthe interface [45]. In the case of DNA detection, the immobilizationof the ssDNA probe molecule (highly charged molecules) leads toan increase in surface capacitance but has an opposite effect uponthe hybridization of ssDNA target molecules due to a removal fromthe surface of the ions contained in the Helmholtz plane (withinof the Debye Length from the electrode), which are contributingto the double layering effect [46]. Although this is true in the caseof DNA short oligos; for antibodies (to see Figs. 1 and 10 in our
Scheme 1. The SPE surface was modified with GNP by electrodeposition to form a nanosized colloidal Au base matrix that produces enhanced surface for antibody immo-bilization. The immobilized monoclonal antibody layer was passivated with a BSA solution that serves as blocking agent. The binding of cTnI in PBS buffer (pH 7.4) leadsto surface charging of electrode surface that is detected by the capacitive component of the impedance extrapolated by the Randles Model. Inset of scheme: parallel platedrawings show capacitive measurements principle. Proteins present a complex charge distribution. Therefore the equivalent capacitance of the electrode/solution interfacechanges when the probe antibodies are immobilized on the GNP surface. The capacitance also changes when cTnI (highly charged molecules) are specifically recognized bymonoclonal antibodies.
766 V. Bhalla et al. / Sensors and Actuators B 161 (2012) 761– 768
Fig. 6. Capacitance based detection of cTnI as compared to sandwich ELISA tests.For ELISA tests, the formation of colored product by alkaline phosphatase reactionwith para-nitrophenyl phosphate substrate was measured at 405 nm. Inset showsthe results of the electrochemical assay after addition of cTnI onto a GNP surfaceimmobilized with a nonspecific antibody.
previous paper in Ref. [44]) this ions removal effect does not hap-pens because of the large size of the proteins. So, the contribution bythe charges carried by the antibody is dominant in case of proteins,as confirmed by the values of Cdl in Table 1.
3.3. Detection of cTnI
The colloidal Au matrix was studied for sensor-to-sensor vari-ation. The GNP deposited sensors show a thin variation from onesensor to another of less than 2% of the average measured capac-itance (data shown in supplementary information Fig. S5). Afterantibody immobilization the GNP antibody surface was passivatedwith BSA solution (1 mg mL−1 in PBS buffer) for 10 min to avoid anynon specific adsorption of proteins [47]. Because of its smaller size,BSA will occupy any leftover regions or crevices in GNP-mAb layer,and will help to protect the surface from ions in solution that mayotherwise directly electrically discharge on the electrodes [44]. BSAalso has a stabilizing effect on the antibody immobilized on thesensor surface [48]. This incubation does not lead to any drasticchange in Cdl, although the increased surface insulation is reflectedby Rct improves upon treatment (Table 1). The BSA-modified SPEwere washed with PBS-T buffer and incubated with antigen (cTnI)followed by washing with PBS-T and impedance scan.
The measured capacitance, with and without cTnI introduction,was quite stable. Data for time stability of capacitance on mAb/BSAwith colloidal Au matrix are quite good (see Fig. S6 in supplemen-tary information). To evaluate the reaction between monoclonalantibody and the antigen, the antibody immobilized GNP/SPE weresubjected to various concentrations of cTnI. The uptake of cTnI leadsto a further increase in Cdl up to a difference of 54.8 nF in case of12 ng mL−1 concentration of the cTnI. An answer to this observa-tion might lie in the highly charged nature of cTnI, its slender shape,and its size of 23 876 Da. It has isoelectric point (pI) of 9.1 and thusit has a net positive charge in PBS buffer at pH 7.4 which leadsfor contributing to Cdl. Fig. 6 compares data for different TnI con-centrations (on a logarithmic scale) with the normalized results.The Cdl based biosensing on cTnI provides a better detection asthe curve shifts towards lower concentrations. The limit of detec-tion [49] is 0.2 ng mL−1 for capacitance based detection, which ismuch better than that of 4.3 ng mL−1 registered with ELISA tests.The upper limit of cTnI in blood for normal people is 0.2 ng mL−1,while values larger than 0.4 ng mL−1 indicate cardiac muscle tis-sue injury or the clear onset of myocardial infarction [50,51]. Thethreshold of 0.4 ng mL−1 is out-of-range of the ELISA tests, while itis well inside the range of the new proposed method. In fact, the
dynamic range for the ELISA test is wider (4.3–100 ng mL−1). Indeedthe Cdl variation ranges from 0.2 ng mL−1 up to 12.5 ng mL−1. Thus,the Cdl based detection has a detection limit which just coincidesthe normal upper limit of cTnI in serum samples and it provides asuitable method for diagnostics of heart failures. The observed lin-earity of the response (R2) over the entire dynamic range is betterfor ELISA (0.99) in comparison to capacitance detection (0.94) evenif considered range of concentrations for the former is larger. Therepeatability coefficient of variation of the sensor (three measure-ments) was ∼5% for 0.4 ng mL−1 concentration of cTnI detection.The total change for the resistive component at a concentration of12 ng mL−1 cTnI was ∼16 k�. The linear fitting of the resistance vs.concentration curve yielded an R2 value equal to 0.81 (Table 1).
In inset to Fig. 6 is shown the change in signal due to thenon-specific adsorption of troponin I antigen. The experiment wasperformed by using the cTnI antigen and by immobilizing theantibody for a different antigen than cTnI. In this case, only non-specific interactions are possible between the cTnI and our colloidalAu modified surfaces. If monoclonal antibody against glycatedhemoglobin (HbA1c) is immobilized on the sensor surface and thesensor is incubated with cTnI antigen, then insignificant changes inimpedance signal are observed. The regeneration of the surface wasalso verified by using 50 mM glycine HCl, pH 2.8, which is knownto dissociate antigen–antibody interactions. The electrodes wereneutralized immediately with 100 mM Tris, at pH 10 to restoreimmobilized antibody. The value of the Cdl registered on regen-erated surfaces was close to that of the freshly prepared colloidalAu surfaces although we observed a small decrease in the resistivecomponent from lower to higher antigen concentrations indicatingthat some weakly bound antibody was removed during the regen-eration process. Thus the approach is good for disposable one-shotuse sensors as well as for reusable-strips application because theCdl is the main suggested parameter for sensing.
4. Conclusions
A capacitance based detection method is presented for theselective detection of the cardiac troponin cTnI. The method pro-vides better sensitivity than usually obtained by conventionalapproaches with sandwich ELISA tests. A new one-step electro-chemical deposition technique is demonstrated for colloidal Aunanostructuration of screen printed electrodes. The formation ofnanosized colloidal Au matrix on SPE was characterized by CVscans and impedance spectroscopy. The successful surface modi-fication was investigated with atomic force microscopy techniqueas well as with electron microscopy. The monoclonal antibodieswere immobilized on the above formed Au matrix and the obtainedbiointerface was characterized by using confocal microscopy. Thebinding reaction of cTnI with anti-cTnI (monoclonal antibodies)was monitored in a label-free manner that showed a positivechange in surface capacitance with increasing of the antigen con-centration. cTnI was successfully detected with a limit of detectionof 0.2 ng mL−1 and showing a good potential to diagnosis of heartfailures with cardiac troponin measure in the blood serum.
Acknowledgements
VB thanks Council of Scientific and Industrial Research (CSIR)for Research Associateship (Scientist pool Scheme). Authors greatlyacknowledge Mr. Deepak at IMTECH for help in acquisition of con-focal microscopy images. Thanks to Dr. Lalit M. Bharadwaj andMr. Deepak at CSIO for their help in SEM characterizations. Mr.H. S. Manjunath is acknowledged for fruitful discussions. AndreaCavallini at EPFL is acknowledged for the English revision of thismanuscript.
V. Bhalla et al. / Sensors and Actuators B 161 (2012) 761– 768 767
Appendix A. Supplementary data
Supplementary data associated with this article can be found, inthe online version, at doi:10.1016/j.snb.2011.11.029.
References
[1] C. Stagni, C. Guiducci, L. Benini, B. Ricco, S. Carrara, C. Paulus, M. Schienle, R.Thewes, A fully electronic label-free DNA sensor chip, IEEE Sensor J. 7 (2007)577–585.
[3] S.D. Bolis, P.C. Charalambous, C.E. Efstathiou, A.G. Mantzila, C.A. Malamou, M.I.Prodromidis, Monitoring of the avidin–biotylinated dextran interaction on Au-and Ti/TiO2 electrode surfaces using a charge integrating device, Sens. Actuat.B: Chem. 114 (2006) 47–57.
[4] E. Antman, J.P. Bassand, W. Klein, M. Ohman, J.L.L. Sendon, L. Rydén, M. Simoons,M. Tendera, Myocardial infarction redefined – A consensus document of TheJoint European Society of Cardiology/American College of Cardiology commit-tee for the redefinition of myocardial infarction: The Joint European Society ofCardiology/American College of Cardiology Committee, J. Am. Coll. Cardiol. 36(2000) 959–969.
[5] M.K. Nusier, B.M. Ababneh Diagnostic efficiency of creatinine kinase (CK),CKMB, troponin T and troponin I in patients with suspected acute myocardialinfarction, J. Health Sci. 52 (2006) 180–185.
[6] L. Wendy, D.J. Cook, L.E. Griffith, M.A. Crowther, P.J. Devereaux, Elevated cardiactroponin levels in critically ill patients: prevalence incidence and outcomes,Am. J. Crit. Care 15 (2006) 280–288.
[7] P.A. Heidenreich, T. Alloggiamento, K. Melsop, K.M. McDonald, A.S. Go, M.A.Hlatky, The prognostic value of troponin in patients with non-ST elevationacute coronary syndromes: a meta-analysis, J. Am. Coll. Cardiol. 38 (2001)478–485.
[8] J. Wei, Y. Mu, D. Song, X. Fang, X. Liu, L. Bu, H. Zhang, G. Zhang, J. Ding, W.Wang, Q. Jin, G. Luoa, A novel sandwich immunosensing method for measuringcardiac troponin I in sera, Anal. Biochem. 321 (2003) 209–216.
[9] Y. Lin, Q. Fu, J. Zhu, J.M. Miller, J. EVan Eyk, Development of a qualitativesequential immunoassay for characterizing the intrinsic properties of circu-lating cardiac troponin I, Clin. Chem. 56 (2010) 1307–1319.
[10] H.C. Yoon, H. Yang, S.Y. Byun, Ferritin immunosensing on microfabricated elec-trodes based on the integration of immunoprecipitation and electrochemicalsignaling reactions, Anal. Sci. 20 (2004) 1249–1253.
[11] H. Chen, C. Jiang, C. Yu, S. Zhang, B. Liu, J. Kong, Protein chips and nanomaterialsfor application in tumor marker immunoassays, Biosens. Bioelectron. 24 (2009)399–3411.
[12] E. Baldrich, F.J. Campo, F.X. Munoz, Biosensing at disk microelectrode arrays.Inter-electrode functionalization allows formatting into miniaturized sensingplatforms of enhanced sensitivity, Biosens. Bioelectron. 25 (2009) 920–926.
[13] J. Lahiri, L. Isaacs, J. Tien, G.M. Whitesides, A strategy for the generation ofsurfaces presenting ligands for studies of binding based on an active ester as acommon reactive intermediate, Anal. Chem. 71 (1999) 777–790.
[14] Y. Wang, R. Wang, Y. Li, B. Srinivasan, S. Tung, H. Wang, M.F. Slavik,C.L. Griffis, Detection of Escherichia coli O157:H7 using interdigitated arraymicroelectrode-based immunosensor, Biol. Eng. Trans. 2 (2010) 49–62.
[15] K. Saha, F. Bender, E. Gizeli, Comparative study of IgG binding to Proteins Gand A: nonequilibrium kinetic and binding constant determination with theacoustic waveguide device, Anal. Chem. 75 (2003) 835–842.
[16] K.V. Singh, D.K. Bhura, G. Nandamuri, A.M. Whited, D. Evans, J. King, R. Solanki,Nanoparticle-enhanced sensitivity of a nanogap-interdigitated electrode arrayimpedimetric biosensor, Langmuir (2011), dx.doi.org/10.1021/la202546a.
[17] E. Hutter, M.P. Pileni, Detection of DNA hybridization by gold nanoparticleenhanced transmission surface plasmon resonance spectroscopy, J. Phys. Chem.B 107 (2003) 6497–6499.
[18] I.D. Vlaminck, P.V. Dorpe, L. Lagae, G. Borghs, Local electrical detection of singlenanoparticle plasmon resonance, Nano Lett. 7 (2007) 703–706.
[19] V.V. Shumyantseva, S. Carrara, V. Bavastrello, J.D. Riley, T.V. Bulko, S.I. Usanov,A.I. Archakov, C. Nicolini, Direct electron transfer between cytochrome P450sccand screen-printed rhodium graphite electrode modified with gold nanoparti-cles, Biosens. Bioelectron. 21 (2005) 217–222.
[20] J.J. Tsai, I.J. Bau, H.T. Chen, Y.T. Lin, G.J. Wang, A novel nanostructured biosensorfor the detection of the dust mite antigen Der p2, Int. J. Nanomed. 6 (2011)1201–1208.
[21] A.C. Cruickshank, A.J. Downard, Electrochemical stability of citrate-capped goldnanoparticles electrostatically assembled on amine-modified glassy carbon,Electrochim. Acta 54 (2009) 5566–5570.
[22] D.R. Reyes, G.I. Mijares, B. Nablo, K.A. Briggman, M. Gaitan, Trapping and releaseof citrate-capped gold nanoparticles, Appl. Surf. Sci. 257 (2011) 8373–8377.
[23] G. Frens, Controlled nucleation for the regulation of the particle size inmonodisperse gold suspensions, Nature (Lond.) Phys. Sci. 241 (1973) 20–22.
[24] P. Sharma, K. Sablok, V. Bhalla, C.R. Suri, A novel disposable electrochemi-cal immunosensor for phenyl urea herbicide diuron, Biosens. Bioelectron. 26(2011) 4209–4212.
[25] S. Eustis, M.A. El-Sayed, Why gold nanoparticles are more precious than prettygold: noble metal surface plasmon resonance and its enhancement of the
radiative and nonradiative properties of nanocrystals of different shapes, Chem.Soc. Rev. 35 (2006) 209–217.
[26] K.V. Singh, A.M. Whited, Y. Ragineni, T.W. Barrett, J. King, R. Solanki, 3D nanogapinterdigitated electrode array biosensors, Anal. Bioanal. Chem. 397 (2010)1493–1502.
[27] N. Wangoo, J. Kaushal, K.K. Bhasin, K.K. Mehta, C.R. Suri, Zeta potential basedcolorimetric immunoassay for the direct detection of diabetic marker HbA1cusing gold nanoprobes, Chem. Commun. 46 (2010) 5755–5757.
[28] H. Huang, Z. Liu, X. Yang, Application of electrochemical impedance spec-troscopy for monitoring allergen–antibody reactions using gold nanoparticle-based biomolecular immobilization method, Anal. Biochem. 356 (2006)208–214.
[29] H. Huang, P. Ran, Z. Liu, Impedance sensing of allergen–antibody interaction onglassy carbon electrode modified by gold electrodeposition, Bioelectrochem-istry 70 (2007) 257–262.
[30] F. Chen, X. Li, J. Hihath, Z. Huang, N.J. Tao, Effect of anchoring groups on single-molecule conductance: comparative study of thiol-amine- and carboxylic-acid-terminated molecules, J. Am. Chem. Soc. 128 (2006) 15874–15881.
[31] L. Venkataraman, J.E. Klare, I.W. Tam, C. Nuckolls, M.S. Hybertsen, M.L. Steiger-wald, Single-molecule circuits with well-defined molecular conductance, NanoLett. 6 (2006) 458–462.
[32] S.Y. Quek, L. Venkataraman, H.J. Choi, S.G. Louie, M.S. Hybertsen, J.B. Neaton,Amine-gold linked single-molecule circuits: experiment and theory, Nano Lett.7 (2007) 3477–3482.
[33] E. Prodan, R. Car, Tunneling conductance of amine-linked alkyl chains, NanoLett. 8 (2008) 1771–1777.
[34] I. Szymanska, H. Radecka, J. Radecki, R. Kaliszan, Electrochemical impedancespectroscopy for study of amyloid �-peptide interactions with (−)nicotine ditartrate and (−) cotinine, Biosens. Bioelectron. 22 (2007) 1955–1960.
[35] G.M. Polya, A.T. Jagendorf, Inactivation of energy linked functions of chloro-plasts at low pH, Arch. Biochem. Biophys. 138 (1970) 540–550.
[36] D.P. Molden, N.M. Alexander, W.E. Neeley, Fetal haemoglobin: optimum con-ditions for its estimation by alkali denaturation, Am. J. Clin. Pathol. 77 (1982)568–572.
[38] F. Patolsky, M. Zayats, B. Katz, I. Villner, Precipitation of an insolubleproduct on enzyme monolayer electrodes for biosensor applications: char-acterization by Faradaic impedance spectroscopy cyclic voltammetry andmicrogravimetric quartz crystal microbalance analyses, Anal. Chem. 71 (1999)3171–3180.
[39] Y. Hou, S. Helali, A. Zhang, N. Jaffrezic-Renault, C. Martelet, J. Minic, T.Gorojankina, M.A. Persuy, E.P. Augy, R. Salesse, F. Bessueille, J. Samitier,A. Errachid, V. Akimov, L. Reggiani, C. Pennetta, E. Alfinito, Immobiliza-tion of rhodopsin on a self-assembled multilayer and its specific detectionby electrochemical impedance spectroscopy, Biosens. Bioelectron. 21 (2006)1393–1402.
[40] A.J. Bard, L.R. Faulkner, Electrochemical Methods: Fundamentals and Applica-tions, second ed., John Wiley, New York, 2000.
[41] S. Carrara, V. Bavastrello, D. Ricci, E. Stura, C. Nicolini, Improvednanocomposite materials for biosensor applications investigated by elec-trochemical impedance spectroscopy, Sens. Actuat. B: Chem. 109 (2005)221–226.
[42] S. Carrara, L. Benini, V. Bhalla, C. Stagni, A. Ferretti, A. Cavallini, B. Riccò, B.Samorì, New insights for using self-assembly materials to improve the detec-tion stability in label-free DNA-chip and immuno-sensors, Biosens. Bioelectron.24 (2009) 3425–3429.
[43] L. Yang, Y. Li, AFM and impedance spectroscopy characterization of the immo-bilization of antibodies on indium-tin oxide electrode through self-assembledmonolayer of epoxysilane and their capture of Escherichia coli O157:H7,Biosens. Bioelectron. 20 (2005) 1407–1416.
[44] S. Carrara, V. Bhalla, C. Stagni, L. Benini, A. Ferretti, F. Valle, A. Gallotta, B. Riccò,B. Samorì, Label-free cancer markers detection by capacitance biochip, Sens.Actuat. B: Chem. 136 (2009) 163–172.
[45] C. Berggren, P. Stalhandske, J. Brundell, G. Johnansson, A feasibility study of acapacitive biosensor for direct detection of DNA hybridization, Electroanalysis11 (1999) 156–160.
[46] C. Guiducci, C. Stagni, G. Zuccheri, A. Bogliolo, L. Benini, B. Samorì, B. Riccò,DNA detection by integrable electronics, Biosens. Bioelectron. 19 (2004)781–787.
[47] R. Singh, I.I. Suni, Minimizing nonspecific adsorption in protein biosensorsthat utilize electrochemical impedance spectroscopy, J. Electrochem. Soc. 157(2010) J334–J337.
[48] J.W. Park, S. Kurosawa, H. Aizawa, S. Wakida, S. Yamada, K. Ishihara, Com-parison of stabilizing effect of stabilizers for immobilized antibodies on QCMimmunosensors, Sens. Actuat. B: Chem. 91 (2003) 158–162.
[49] S. Carrara, A. Cavallini, V. Erokhin, G.D. Micheli, Multi-panel drugs detectionin human serum for personalized therapy, Biosens. Bioelectron. 26 (2011)3914–3919.
[50] R.P. Relos, I.K. Hasinoff, G.J. Beilman, Moderately elevated serum tro-ponin concentrations are associated with increased morbidity and mortalityrates in surgical intensive care unit patients, Crit. Care Med. 31 (2003)2598–2603.
[51] S.E. Kay, J. Doery, D. Sholl, Clozapine associated pericarditis and elevated tro-ponin I, Aust. N.Z. J. Psychiatry 36 (2002) 143–144.
768 V. Bhalla et al. / Sensors and Actuators B 161 (2012) 761– 768
Biographies
Vijayender Bhalla holds a Masters in Medical Biochemistry from Postgraduate Insti-tute of Medical Education and Research (PGIMER), got in Chandigarh, India. Heobtained his Ph.D. in Bionanotechnology jointly from CSIO(CSIR) and Panjab Uni-versity in Chandigarh, India, in 2006. After graduating, Vijayender joined as postdocthe group of Prof. Bruno Samori at the University of Bologna, in Italy. In bologna, hisresearch was focused on cancer markers detection based on capacitance biochip.Later on he moved to Montreal to join Dr. Valter Zazubovich at Concordia Uni-versity, Canada. His research at Concordia was mainly focused on photosyntheticbiosensors for detection of herbicides and explosives. Now, Dr. Bhalla is a SRA (CSIR,Pool Scientist) at Biosensors Laboratory, Microbial Research Institute in Chandigarh.His research is mainly focused on development of label-free immunosensors forenvironment and healthcare applications.
Sandro Carrara is Senior Scientist and Lecturer at the École Polytechnique Fédéralede Lausanne, and Professor of Optical and Electrical Biosensors at the Departmentof Electrical Engineering and Biophysics (DIBE) of the University of Genoa. He wasProfessor of Biophysics at the University of Genoa and Professor of Nanobiotech-nology at the University of Bologna. His main scientific interest is on electricalphenomena mediated by nano-structured molecular thin films and the develop-ment of bio/nano/chips. He published more than 100 scientific papers and heobtained 10 industrial patents. His work has been awarded several times (Nan-otera 2011, in Bern, NanoEurope 2009 in Rapperswil, IEEE Conference PRIME2010 in Berlin, and 2009 in Cork, NATO Advanced Research 1996 in Sezged, best
referees’ of Biosensor and Bioelectronics Journal in 2007). Sandro Carrara is a Founderand Editor-in-Chief of the journal BioNanoScience, and an Associate Editor of theIEEE Sensors Journals and of the IEEE Transactions on Biomedical Circuits and Sys-tems. He is/was in the boards of several IEEE International Conferences such asIWASI/Bari-2011, ISMICT/Montreux-2011, BioCAS/San Diego 2011, BioCAS/Cyprus2010, BioCAS/Beijing 2009, NanoNets/Luzern 2009.
Priyanka Sharma got her Masters in Environmental Sciences at the University ofPanjab, Chandigarh, India. She is presently working as Sr. research fellow (UGC) forher PhD with Dr. C. Raman Suri at the IMTECH Institute, in Chandigarh, India. Shehas also worked as a visiting research scholar at Northwestern University (USA)for 4months. Her research is focused on molecular receptors and assay development fordetection of phenyl urea herbicides. She has 4 international papers on the detectionof Diuron.
Yogesh Nangia did his Masters in Biotechnology from Bhopal, India. Currently heis working as a senior research fellow (CSIR). He has three years experience inbiological systems based diagnostics and synthesis of molecular receptors such asaptamers, ScFv and antibodies against small molecules and proteins.
C. Raman Suri received his Master Degree in Electronics from AMU, Aligarh andPh.D. in Biotechnology from Punjab University, Chandigarh, India, specializingin immunosensor development. Currently he is working as a Senior Scientist atInstitute of Microbial Technology, Chandigarh, India. His current activities includehapten-protein conjugation for antibodies generation, immunoassay developmentand sensor design.
