Piezomicrogravimetric and Impedimetric Oligonucleotide Biosensors

Jul 5, 2013 - Chemistry Department, Central University of Rajasthan, Kishangarh, Rajasthan, India. ∥. Faculty of Mathematics and Natural Sciences, S...
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Piezomicrogravimetric and Impedimetric Oligonucleotide Biosensors Using Conducting Polymers of Biotinylated Bis(2,2′-bithien-5yl)methane as Recognition Units Marta Sosnowska,†,‡ Piotr Pieta,† Piyush S. Sharma,† Raghu Chitta,§ Chandra B. KC,‡ Venugopal Bandi,‡ Francis D’Souza,*,‡ and Wlodzimierz Kutner*,†,∥ †

Department of Physical Chemistry of Supramolecular Complexes, Institute of Physical Chemistry, Polish Academy of Sciences, Kasprzaka 44/52, 01-224 Warsaw, Poland ‡ Department of Chemistry, University of North Texas, 1155 Union Circle No. 305070, Denton, Texas 76203-5017, United States § Chemistry Department, Central University of Rajasthan, Kishangarh, Rajasthan, India ∥ Faculty of Mathematics and Natural Sciences, School of Science, Cardinal Stefan Wyszynski University in Warsaw, Wóycickiego 1/3, 01-815 Warsaw, Poland S Supporting Information *

ABSTRACT: A new conducting polymer of biotinylated bis(2,2′-bithien-5-yl)methane was prepared and applied as the recognition unit of two different biosensors for selective oligonucleotide determination using either electrochemical impedance spectroscopy (EIS) or piezoelectric microgravimetry (PM) for label-free analytical signal transduction. For preparation of this unit, first, a biotinylated bis(2,2′-bithien-5yl)methane functional monomer was designed and synthesized. Then, this monomer was potentiodynamically polymerized to form films on the surface of a glassy carbon electrode (GCE) and a Au electrode of a quartz crystal resonator (QCR) for the EIS and PM transduction, respectively. On top of these films, neutravidin was irreversibly immobilized by complexing the biotin moieties of the polymer. Finally, recognizing biotinylated oligonucleotide was attached by complexing the surface-immobilized neutravidin. This layer-by-layer assembling of the poly(thiophene−biotin)−neutravidin− (biotin−oligonucleotide) recognition film served to determine the target oligonucleotide via complementary nucleobase pairing. Under optimized determination conditions, the target oligonucleotide limit of detection (LOD) was 0.5 pM and 50 nM for the EIS and PM transduction, respectively. The sensor response to the target oligonucleotide was linear with respect to logarithm of the target oligonucleotide concentration in a wide range of 0.5 pM to 30 μM and with respect to its concentration in the range of 50 to 600 nM for the EIS and PM transduction, respectively. The biosensors were appreciably selective with respect to the nucleobase mismatched oligonucleotides.

T

polymerization resulting in deposition of a conducting polymer bearing an intact biotin moiety available for subsequent avidin binding. This biotin−avidin immobilization strategy eliminates the need of using additional reagents adsorbed on surface of the polymer film for this immobilization. Similar works reporting on formation of a conducting polymer bearing biotin on its surface were reported for terthiophenes7,8 and polypyrrole.1,9,10 However, the pyrrole-based monomers are stable only in their doped state,11 which significantly limits the scope of their application. Moreover, the synthetic route of preparation of a monomer for the terthiophene-based polymers was much more

he biotin−avidin binding is a widely employed mechanism for biosensor development.1−5 The apparent reason for this is the very high stability constant of the resulting complex being of the order of 1015 M−1.6 This high stability allows for easy formation of stable assemblies of substances, one of which being derivatized with biotin and the other with avidin, even though only minute quantities of the derivatized substances are available. Moreover, biotinylated reagents, like enzymes, proteins, antibodies, and oligonucleotides, as well as kits allowing for biotinylation of different compounds with high yield in the user lab, are becoming readily commercially available. Herein, we report on application of a newly prepared functional monomer of bis(2,2′-bithien-5-yl)methane derivatized with biotin (Scheme 1) for oligonucleotide sensing. The presence of the bis(2,2′-bithien-5-yl)methane moeity in the monomer allows for straightforward oxidative electrochemical © XXXX American Chemical Society

Received: May 9, 2013 Accepted: July 5, 2013

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EXPERIMENTAL SECTION Chemicals. 2,2′-Bithiophene, 4-hydroxy-benzaldehyde, 70% HClO 4 , Na 2 HPO 4 , KH 2 PO 4 ·3H 2 O, KCl, 1-ethyl-3-(3dimethylaminopropyl)carbodiimide (EDCI), and K3Fe(CN)6 were purchased from Sigma-Aldrich. D-Biotin and neutravidin were supplied by ChemPep and Thermo Scientific, respectively. NaOH, NaCl, K4Fe(CN)6, and acetonitrile were purchased from Fisher. Tetra-n-butylammonium perchlorate, (TBA)ClO4, and ethanol were from Southwestern Analytical, Inc., and PharmacoAapek, respectively. All chemicals were used as received. The 21base oligonucleotides were purchased from Alpha DNA. Sequences of the oligonucleotides were as follows: recognizing biotinylated oligonucleotide 5′-ATG TGG AAA ATC TCT AGC AGT-3′, complementary target oligonucleotide 5′-ACT GCT AGA GAT TTT CCA CAT-3′, seven-base mismatched target oligonucleotide 5′-ACG GCG TGA GAA TGT CTA CGT-3′, and two-base mismatched target oligonucleotide 5′ACA GCA AGA GAT TTT CCA CAT-3′. All oligonucleotide solutions were prepared immediately before measurements using 0.1 M phosphate buffer saline (PBS) of pH = 7.4. Synthetic procedure of bis(2,2′-bithien-5-yl)-(4hydroxyphenyl)methane biotin ester preparation consisted of two steps. First, bis(2,2′-bithienyl)-(4-hydroxyphenyl)methane was prepared using previously described procedure.37 Then the product was used for esterification with biotin in the presence of EDCI as a carboxyl activating agent. The synthetic procedure is described into more detail in the Supporting Information. Instrumentation and Procedures. Potentiodynamic, Cyclic Voltammetry, and Electrochemical Impedance Spectroscopy Measurements. These measurements were performed with the computerized PARSTAT 2273 potentiostat/galvanostat controlled by the PowerSuite software, both of Princeton Applied Research or an AUTOLAB computerized electrochemistry system of Eco Chemie equipped with expansion cards of the PGSTAT 12 potentiostat and the FRA2 frequency response analyzer and controlled by the GPES 4.9 software of the same manufacturer. A Pt plate, Ag|AgCl, and 3 mm diameter glassy carbon disk or a gold film coated glass slides (7 × 7 mm2) served as the auxiliary, reference, and working electrode, respectively. The EIS spectra were recorded in the 100 kHz to 500 MHz frequency range for GCE immersed in 0.1 M PBS, which was 0.1 M in K3Fe(CN)6 and 0.1 M in K4Fe(CN)6. For a given target oligonucleotide, first, an open-circuit potential was measured. Then, the EIS experiments were performed at just this potential being applied. Experimental data from the EIS measurements were fitted with parameters of equivalent circuits using ZView software of Scribner Associates, Inc. Piezoelectric Microgravimetry Experiments. These experiments were performed using the electrochemical quartz crystal microbalance model EQCM 5710 of IPC PAS under control of the EQCM 5710-S2 software of the same manufacturer. The ATcut, plano−plano, 10 MHz QCRs of 14 mm in diameter, coated on both sides with electrodes of 5 mm diameter 100 nm thick Au films evaporated over Ti underlayer films, were used. Prior to the polymer film electrodeposition, the resonators were cleaned with the “piranha” solution (H2O2/H2SO4, 1:3, v/v for 10 s (caution: the “piranha” solution is dangerous when contacting skin or eye as it violently reacts with most organic compounds). A Pt coil and a AgCl film coated Ag wire were used as the counter and pseudoreference electrode, respectively. For electropolymerization, which was simultaneously performed under PM and potentiodynamic batch-solution conditions, the EQCM 5710

Scheme 1. Structural Formula of the Functional Monomer of Bis(2,2′-bithien-5-yl)-(4-hydroxyphenyl)methane Biotin Ester

complicated and tedious than that developed in the present work. That would significantly increase cost of a sensor based on the terthiophene monomer. Surprisingly, the terthiophene monomers were used to prepare sensors only for avidin thus limiting their possible use in constructing devices of significance to analytical chemistry. Herein, we provide a complete test of the designed and prepared biotinylated monomer with respect to its application for chemosensor preparation. In order to demonstrate its versatility, we used two different transduction techniques and applied them for the oligonucleotide determination. The importance of the DNA determination using biosensors is manifested by its broad application ranging from forensic uses and medical diagnostics to detection of toxins, whole bacteria, and viruses.10,12−16 Many different transduction platforms have been used for this sensing including optical,17 surface plasmon resonance,3 and fluorescence18,19 spectroscopy as well as those based on detection of products of enzymatic reactions,2,20 redox reactions of ferrocene modified oligonucleotides,21,22 or direct detection of the DNA intercalation probes.23−25 Other platforms involve piezoelectric microgravimetry (PM) using quartz crystal resonators (QCRs)1,12,26 or microcantilevers,27 as well as the field effect transistors.28,29 Moreover, DNA sensors frequently use electrochemical impedance spectroscopy (EIS) for highsensitivity signal transduction.13,30−35 Herein, two label-free transductions, viz., PM and EIS, were exploited for development of two different oligonucleotide biosensors. In order to prepare the biosensor recognition unit, first, the herein synthesized functional monomer was electropolymerized to result in a thin conducting polymer film on the surface of either a QCR electrode or glassy carbon electrode (GCE) transducer. Then, this film was irreversibly decorated with neutravidin. This step allowed for subsequent immobilization of a recognizing biotinylated oligonucleotide. The layer-bylayer obtained recognition film was used for determination of a complementary oligonucleotide. For the present study, we selected an oligonucleotide of the 5′-ACT GCT AGA GAT TTT CCA CAT-3′ sequence, which was reported as the singlestranded DNA fragment of the HIV-1 virus.36 The detection signal of the change of the resonant frequency of QCR in PM under the flow injection analysis (FIA) conditions and that of the film resistance in EIS under batch analysis conditions was due to hybridization of the polymer-immobilized recognizing oligonucleotide and the target complementary nucleotide present in solution. B

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1.30 V (Figure 1a) resulted from oxidation of the bisbithiophene moiety of the functional monomer leading to formation of the polymer film. This film formation is confirmed with the use of EQCM by the frequency decrease (curve 1 in Figure 1b) indicating the mass increase (curve 2 in Figure 1b) of the film deposited on the surface of the QCR, as recalculated using the Sauerbrey equation. A positive potential shift of the bisbithiophene moiety oxidation can be attributed to the increase of the electrode charge-transfer resistance due to the growth of the film much less conductive than the bare electrode. Similar behavior has already been reported for polymerization of other monomers of the bis(2,2′-bithien-5-yl)methane polymerizing moiety.39,40 Supporting Information Figure S1 shows AFM images of the polymer film deposited during one (Supporting Information Figure S1a), three (Supporting Information Figure S1b), five (Supporting Information Figure S1c), and seven current− potential (I−E) cycles (Supporting Information Figure S1d). The film deposited during one I−E cycle was uniform, covering the entire surface of the Au-coated glass slide (Supporting Information Figure S1a). The film consisted of merged polymer grains of 12 ± 2 nm in diameter (Supporting Information Figure S1b). With the increase of the number of I−E cycles, the film thickness and roughness increased. Apparently, the film became more porous with holes, indicated by arrows in Supporting Information Figure S1, parts c and d. The rate of the film growth was lower the higher was the number of the I−E cycle, in line with the PM measurements. That means that the mass increase was smaller for each consecutive cycle. Scheme 2 shows the AFM images of the biosensor recognition film taken after each step of its preparation: the polymer film deposited using three I−E cycles (layer a and image a′), the polymer film with immobilized neutravidin (layer b and image b′), then with the recognizing oligonucleotide (layer c and image c′), and after binding the target oligonucleotide (layer d and image d′). Because of the film modification with neutravidin, morphology of the film distinctly changed. That is small, 12 ± 2 nm in diameter, grains (Scheme 2a′) were replaced by clusters of 25 ± 1 nm in diameter (Scheme 2b′). This grain aggregation may indirectly suggest that neutravidin was complexed by biotin. The consecutive immobilization of recognizing oligonucleotide made the film more rigid with well-seen pores of 17 ± 4 nm in diameter (Scheme 2c′). Moreover, flat terraces that appeared may suggest aggregation of the oligonucleotides, indicating that the oligonucleotide structure is responsible for the film solidification. Hybridization of the target oligonucleotide did not change morphology of the film appreciably (Scheme 2d′). Impedance Response of the Polymer-Coated GCE. Figure 2 shows Nyquist plots for GCE at each step of the biosensor preparation, and then after exposing this electrode to the target oligonucleotide. Each curve consists of a flatten semicircle and a straight line in the high- and low-frequency range, respectively, typical for a simple redox process under the charge-transfer rate control of a porous system.41,42 In order to determine the value of the charge-transfer resistance, Rct, the experimental data were fitted with electric parameters of the modified Randles−Ershler circuit (inset in Figure 2). This circuit provided a good fit of the impedance data, in agreement with fits of impedance of other films used for DNA sensing.31,33−35 In the equivalent circuit adopted (inset in Figure 2), the Rs element represents the solution resistance; its value is constant (22.3 Ω, 6.3 Ω cm2 for the bare GCE) for all measured films. For electrodes coated with the polymer films, the double layer capacity, Cdl, used in the Randles circuit, was herein replaced by

microbalance was interfaced with the EP-20 potentiostat of IPC PAS. Flow Injection Analysis Measurements. The QCR, coated with the polymer film, was mounted in the flow-through EQCM 5610 holder38 of IPC PAS to examine analytical performance of the biosensor under FIA conditions. A solution of 0.1 M PBS (pH = 7.4), used as the carrier solution, was pumped through the holder with the effusion syringe pump model KDS100 of KD Scientific, Inc. The FIA experiments were carried out at the 35 μL/min flow rate. The analyte samples, dissolved in PBS of the same composition as that of the carrier solution and the 100 μL volume, were injected with the use of the rotary six-port loop valve model 7725i of Rheodyne. Atomic Force Microscopy Imaging. The biosensor recognition film was imaged by atomic force microscopy (AFM) at each step of its preparation using a MultiMode 8 AFM under control of Nanoscope V controller of Bruker. The ScanAsyst mode was utilized for sample imaging with the use of the ScanAsyst-Air-HR probes provided by Bruker. The films for imaging were deposited on the gold film coated glass slides. The slides were cleaned with acetone and then dried in an Ar stream prior to use. Preparation of the Biosensors. Detailed procedures of the biosensor preparation are described in the Supporting Information.

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RESULTS AND DISCUSSION Deposition of the Biotinylated Polymer. Potentiodynamic deposition of this polymer on an electrode surface was performed by linear cycling the potential in the range of 0.50 to 1.50 V (Figure 1). An anodic peak observed between 1.20 and

Figure 1. (a) Current−potential curve at GCE (b) corresponding resonant frequency (curve 1) and mass changes (curve 2) at QCR in the course of potentiodynamic electropolymerization of the functional monomer in 0.1 M (TBA)ClO4 in acetonitrile/ethanol (1:1, v/v). The potential scan rate was 50 mV/s. C

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Scheme 2. Consecutive Steps of Preparation of the Recognition Film on the Electrode Surface and AFM Images Prepared for (a) Polymer Film Prepared Using Three I−E Cycles on a Gold Film Coated Glass Slide and This Film after Modification with (b) Neutravidin, (c) Recognizing Oligonucleotide, and (d) Target Oligonucleotide

component.43 Equation 1 shows the impedance contribution of CPE, ZCPE, to the total impedance.

ZCPE =

1 (iω)ϕ T

(1)

In this equation, i = (−1)1/2, ω = 2πf, ω is the angular frequency, T is the frequency-independent proportionality factor with physical meaning related to the diffusion coefficient and characteristic of the electrochemical system, and ϕ is an exponent. This exponent can take values between 0 and 1. For ϕ = 1, the system behavior is purely capacitive where impedance of the element, ZCPE, equals to C−1, while for ϕ = 0 ZCPE equals to pure resistance. The Warburg impedance, W, was also introduced to the equivalent circuit (inset in Figure 2). This element represents mass transfer to the electrode accompanying the redox process. Contribution of the Warburg impedance to the total impedance is described by Equation 2, as follows

Figure 2. Exemplary Nyquist plots for GCE (1) coated with the polymer film of the biotinylated monomer, and then after consecutive immobilization of (2) neutravidin, (3) biotinylated oligonucleotide, and (4) after equilibration with 50 pM target oligonucleotide for 6 min. Measurements were performed for 0.1 M K3Fe(CN)6 and 0.1 M K4Fe(CN)6 at an open-circuit potential. The inset shows the modified Randles−Ershler equivalent circuit used to fit experimental data: Rs, solution resistance; Rct, charge-transfer resistance; CPE, constant phase element; W, Warburg impedance.

ZW =

B tanh(iTω) (iTω)ϕ

(2)

where B is the fitted parameter. Herein, Rct describes the charge-transfer resistance of the electron transfer between the electrode and the redox species in solution. Therefore, Rct corresponds to the resistance of the electrode on which a faradaic reaction proceeds. The value of this

the constant phase element (CPE) because the film did not behave as an ideal capacitor and, therefore, had nonzero real D

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value after 10 min immersion of the electrode in the 30 μM recognizing oligonucleotide solution (data not shown). Analytical Performance of the Impedimetric Biosensor. After deposition of a complete recognition film, the electrode was immersed in the target oligonucleotide solutions of different concentrations for 6 min. Then, the electrode was rinsed with 0.1 M PBS (pH = 7.4), and the potential was linearly cycled between −0.20 and 0.60 V. This biosensor treatment resulted in repeatable impedimetric data (Figure S3 in the Supporting Information). Next, the experimental impedance data were fitted with electric parameters of the equivalent circuit presented in the inset of Figure 2, and the determined Rct values were used to construct the calibration plot (curve 1 in Figure 3).

resistance is affected by any surface changes leading to blocking of the electrode, for instance, by deposition of a less conducting film. Here, Rct was calculated for each step of preparation of the recognition film and then after interaction of this film with the target oligonucleotide (Table S1 in the Supporting Information). As expected, both the exponent ϕ and Rct increased after immobilization of each layer of the recognition film. This increase indicates that the film became more insulating and its electrical double layer properties remained as those of an ideal capacitor. Moreover, the T value changed only slightly suggesting that diffusion of the electroactive species was not much affected by modification of the film (Supporting Information Table S1). Our present results are close to those reported for a similar system.35 Optimization of the Recognition Film Preparation for the EIS Biosensor. In order to prepare biosensors of appreciable analytical performance, i.e., low limit of detection (LOD) and high sensitivity and selectivity, several parameters of preparation of the recognition films were optimized. First, influence of the polymer film thickness on sensor performance was examined (Table S2 in the Supporting Information). This thickness was controlled with the number of I−E cycles.44 The AFM imaging showed (Supporting Information Figure S1) that the electrode surfaces were completely coated by the film under each of the deposition conditions. However, deposition of a thicker film resulted in the increase of both its roughness and Rct. It means that the electron transfer through the polymer film between the electrode substrate and the electroactive species in the solution can be significantly hindered due to the increase of the film resistivity. In consequence, the Rct change due to immobilization of neutravidin, and then the recognizing oligonucleotide, would be indiscernible. This behavior was confirmed experimentally. That is, immobilization of neutravidin and the recognizing oligonucleotide caused merely a small Rct increase for thicker films. In effect, the sensitivity to the target oligonucleotide was low. In order to determine optimum film thickness, calibration plots were constructed for the recognition films prepared using different numbers of I−E cycles (Figure S2 in the Supporting Information). Apparently, the optimum number of cycles for the 1 mM monomer was three, providing effective resistivity and surface roughness (or porosity) of the resulting polymer film. Further deposition of the polymer increased porosity of the film on the one hand and increased its resistivity (Table S2 in the Supporting Information) and, in consequence, decreased its LOD on the other. Other parameters affecting the analytical performance of the biosensors include time of the electrode immersion in the neutravidin solution, and then in the recognizing oligonucleotide solution. Seemingly, the larger the amount of the recognizing oligonucleotide immobilized on the electrode, the larger the amount of the target oligonucleotide could be bound. Hence, sensitivity could be higher. Therefore, the dependence of Rct on the immersion time was investigated. For that, the electrode coated with the polymer film was immersed in the 1 mg/mL neutravidin solution for 2, 5, 7, 10, 15, and 20 min. It appeared that the Rct value was higher the longer was the immersion time up to 10 min, and then it remained constant suggesting saturation of the active sites of the polymer responsible for the biotin−neutravidin complexation (data not shown). Similarly, the time of the recognizing oligonucleotide immobilization was optimized. It turned out that the Rct reached its maximum steady

Figure 3. Nyquist plots for the GCE coated by the recognizing film after immersion in solution of (1) the target oligonucleotide, (2) twonucleobase mismatched target oligonucleotide, (3) seven-nucleobase mismatched target oligonucleotide, and (4) recognition film before interaction with the target oligonucleotide; the inset presents calibration plots constructed using the data obtained by fitting electric parameters of the equivalent circuit shown in the inset in Figure 2 to the experimental data. Measurements were performed for 0.1 M K3Fe(CN)6 and 0.1 M K4Fe(CN)6 at the applied potential equal to the opencircuit potential.

The sensor response was linear with respect to the logarithm of the target oligonucleotide concentration in a quite wide concentration range (0.5 pM to 30 μM). The linear regression equation and the correlation coefficient of the calibration plot was (Rct,f − Rct,i)/Ω = (19.91 ± 0.59) + (81.18 ± 1.48) log(ctarget/ nM) and 0.99, respectively. In this equation, ctarget is the concentration of the target oligonucleotide, while Rct,i and Rct,f are the charge-transfer resistances of the initial recognition film and the recognition film after immersion in the target oligonucleotide solution, respectively. The recognition films, prepared under the same conditions, were used to compare the biosensor selectivity with respect to oligonucleotides containing two- and seven-nucleobase mismatches (curves 2 and 3, respectively, in Figure 3). The sensitivity with respect to the target oligonucleotide, 19.91 ± 0.59 Ω, was nearly twice that to the two-nucleobase mismatched oligonucleotide, 10.87 ± 0.56 Ω, and 7 times that to the seven-nucleobase mismatched oligonucleotide, 2.83 ± 1.47 Ω. With respect to its LOD and linear concentration range, the performance of the current impedimetric sensor is comparable to other oligonucleotide sensors using EIS for signal transduction.13,30−35 E

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Analytical Performance of the Piezoelectric Microgravimetry Biosensor. For preparation of the recognition film for the PM biosensor, the same conditions were used as those described above for preparation of the impedimetric biosensor (i.e., three I−E cycle film depositions and 1 mM monomer solution). The recorded under FIA conditions time dependence of the frequency and dynamic resistance changes accompanying the consecutive immobilization steps of the recognition film preparation is presented in Figure 4. Almost every solution

5 shows the time dependence of simultaneously recorded resonant frequency change and the dynamic resistance change

Figure 5. Simultaneous resonant frequency and dynamic resistance changes with time accompanying (a) consecutive injections of 600 nM target oligonucleotide, marked with black arrows, and followed by injections of 0.1 M NaOH for dehybridization, marked with red arrows (b), under flow injection analysis (FIA) conditions. The injected sample volume was 100 μL, the carrier solution was PBS (pH = 7.4), and its flow rate was 35 μL/min.

Figure 4. Changes with time of (a) the resonant frequency and (b) dynamic resistance accompanying consecutive injections of (1) 1 mg/ mL neutravidin, (2) 15 μM biotinylated oligonucleotide, subsequently (3) 15 μM noncomplementary oligonucleotide, and then (4) 15 μM target oligonucleotide under flow injection analysis (FIA) conditions for the recognition polymer film. The injected volume of the sample in PBS (pH = 7.4) was 100 μL; the flow rate of the 0.1 M PBS (pH = 7.4) carrier solution was 35 μL/min.

for five consecutive injections of solutions of 600 nM target oligonucleotide each followed by injection of 0.1 M NaOH. After each oligonucleotide injection, frequency dropped, expectedly, to reach a constant value. Evidently, the recognizing oligonucleotide was permanently hybridized by the target oligonucleotide. Due to the 0.1 M NaOH injection, initially the frequency abruptly dropped, which was mirror imaged by a sudden raise of the dynamic resistance. When the PBS carrier solution reached the flow-through cell, i.e., after ∼3 min after the initial frequency drop, the frequency returned to its initial baseline level. This behavior indicates a complete elution of the target oligonucleotide from the recognition film by 0.1 M NaOH. Presumably, this abrupt frequency drop and subsequent increase to its baseline level was due to a change in swelling of the polymer accompanying switching of the flow of the 0.1 M PBS carrier solution (pH = 7.4) to that of the 0.1 M NaOH solution (pH = 13.0). Signal repeatability was estimated from the data shown in Figure 5. The determined standard deviation (n = 5) of the target oligonucleotide quantization was 13%. The calibration curve (Figure 6) was constructed from the frequency change with time during consecutive injections of the target oligonucleotide of different concentrations, each followed by injection of 0.1 M NaOH. The biosensor response to the target oligonucleotide was linear in the concentration range of 50−600 nM obeying the linear regression equation of Δf/Hz = (0.028 ± 0.003) + (4.74 ± 0.66)ctarget/nM with the correlation coefficient of 0.94 and Δf being the resonant frequency change due to injection of the target oligonucleotide. The herein determined low, 50 nM, LOD corresponding to 3 μg/mL is similar to that of the reported PM sensors for biological species, e.g., proteins, oligonucleotides, which is most typically in the range of micrograms per milliliter.1,12,26,46 Moreover, the value of the stability constant, Ks, of the oligonucleotide duplex formed by hybridization of the immobilized recognizing nucleotide with the target oligonucleotide was determined using the literature procedure.47 This

injection caused a frequency decrease (Figure 4a) corresponding to consecutive irreversible immobilization of a different species. From this decrease, the mass of the immobilized species was calculated according to the Sauerbrey equation, described in the Supporting Information. The frequency decrease was accompanied by a small dynamic resistance increase (Figure 4b) in the range of few ohms, proving that the viscosity of the film did not change much. Apparently, rigidity of the film remains practically unchanged after each step of its preparation. The mass of neutravidin immobilized on the polymer film, calculated from the frequency change shown in Figure 4, was 276 ng (4.6 fmol). The mass of the recognizing oligonucleotide subsequently immobilized was 87 ng (13.5 fmol). Hence, the immobilized neutravidin-to-(recognizing oligonucleotide) mole ratio was 2.9 being close to that of 3, expected. The reason for this is that the neutravidin has four biotin binding sites,45 at least one of which is occupied by the biotin moiety of the polymer film. Hence, the recognition film is fully modified with the recognizing oligonucleotide. Subsequent injection of the target oligonucleotide solution resulted in the mass change of 63 ng (9.8 fmol). Apparently, hybridization of the recognizing oligonucleotide with the target oligonucleotide was incomplete by 3.7 fmol after the first injection. Most likely, that was because of significantly lower stability of the hybridization complex (calculated below) compared to that of the neutravidin−biotin complex. After preparation of the recognition film, the calibration curve was constructed by injecting the target oligonucleotide solutions of different concentrations. After each injection, a sample of the 0.1 M NaOH solution was injected to unzip the double-stranded oligonucleotide formed, thus regenerating the biosensor. Figure F

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meet particular sensing needs by selection of the most appropriate transduction technique. Moreover, there are countless analytes whose determination can be elaborated using the hereby presented methods since the biotin−avidin chemistry and procedures of biotinylation of different species are well-known. Noteworthy, the electropolymerization used for polymer deposition is a straightforward, low cost, and repeatable way of the transducer surface modification, thus making biotin immobilization very attractive.

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ASSOCIATED CONTENT

S Supporting Information *

Preparation of bis(2,2′-bithien-5-yl)-(4-hydroxyphenyl)methane biotin ester, preparation of the biosensors, preparation of samples for AFM imaging, EIS data, and Sauerbrey equation. This material is available free of charge via the Internet at http:// pubs.acs.org.

Figure 6. Resonant frequency change against the target oligonucleotide concentration during flow injection analysis (FIA) PM measurements. After each injection of the target oligonucleotide, 0.1 M NaOH was injected. Volume of the injected sample solution was 100 μL, the carrier solution was 0.1 M PBS (pH = 7.4), and its flow rate was 35 μL/min.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected] (F.D.); [email protected] (W.K.). Phone: +940-369-8832 (F.D.); +48 22 343 3217 (W.K.). Fax: +940 565 4318 (F.D.); +(48 22) 343 3333 (W.K.).

procedure is described in detail in the Supporting Information. The determined complex stability constant was, Ks = 4.56 × 103 M−1.

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Notes

The authors declare no competing financial interest.

CONCLUSIONS Newly synthesized biotinylated bis(2,2′-bithien-5-yl)methane monomer efficiently electropolymerized under potentiodynamic conditions. The biotin-containing resulting polymer film irreversibly bound neutravidin. This binding enabled preparation of a label-free biosensor for sequence-specific oligonucleotide determination. LOD of the devised and herein prepared piezomicrogravimetric biosensor was higher than that of the impedimetric biosensor (50 nM and 0.5 pM, respectively). Advantageously, however, it operated under the FIA conditions and was capable of reversible binding of the target oligonucleotide. The PM and EIS biosensors were appreciably selective with respect to mismatched target oligonucleotides. The difference in LODs of both biosensors prepared mainly arises from technical limitations that are different for each technique. For the EQCM used, the measurable mass change is in a single nanogram range. Therefore, no lower LOD could be reached. These differences in LODs may be used as guidance to allow for selecting the appropriate technique for a given analyte concentration range, which is mainly defined by the sample preparation procedure adopted. Apparently, the synthesized monomer is versatile being suited for preparation of biosensors using different transduction schemes. The EIS results showed that the resistance of the recognition film changed due to binding of the target oligonucleotide thus indicating that other less sophisticated electrochemical techniques, e.g., differential pulse voltammetry (DPV), could be used for transduction. Moreover, the results of the PM experiments proved a significant mass change due to binding of the target oligonucleotide. Hence, other techniques based on the mass change of the surface confined film, e.g., surface plasmon resonance (SPR) spectroscopy, could be applied for the transduction. Justifiably, therefore, it is to claim that the monomer opens a new avenue of biosensor preparation, as the same film preparation procedure can be used for preparation of biosensors of different properties. So, biosensor properties, like detectability and the linear concentration range, can be tuned to

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ACKNOWLEDGMENTS The financial support of the Foundation for Polish Science (MPD/2009/1/styp15) to M.S., the U.S. National Science Foundation (Grant No. 1110942) to F.D., the European Union 7.FP under Grant REGPOT-CT-2011-285949-NOBLESSE to P.P., and the European Union within European Regional Development Fund, through Grant Innovative Economy (POIG.01.01.02-00-008/08) to W.K. is gratefully acknowledged. Access to the AFM instrumentation was funded by the Foundation for Polish Science under the FOCUS Programme No. FG 3/2010 Grant.

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