Electrochemical Label-free and Reagentless Genosensor Based on

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Electrochemical Label-free and Reagentless Genosensor Based on an Ion Barrier Switch-off System for DNA Sequence-Specific Detection of the Avian Influenza Virus Katarzyna Kurzątkowska,† Agnieszka Sirko,‡ Włodzimierz Zagórski-Ostoja,‡ Wim Dehaen,§ Hanna Radecka,† and Jerzy Radecki*,† †

Institute of Animal Reproduction and Food Research, Polish Academy of Sciences, Tuwima 10, 10-748 Olsztyn, Poland Institute of Biochemistry and Biophysics, Polish Academy of Sciences, Pawińskiego 5A, 02-106 Warsaw, Poland § Department of Chemistry, University of Leuven, Celestijnenlaan 200F, B-3001 Leuven, Belgium ‡

S Supporting Information *

ABSTRACT: This paper concerns the development of genosensors based on redox-active monolayers incorporating (dipyrromethene)2Cu(II) and (dipyrromethene)2Co(II) complexes formed step by step on a gold electrode surface. They were applied for electrochemical determination of oligonucleotide sequences related to avian influenza virus (AIV) type H5N1. A 20mer probe (NH2-NC3) was covalently attached to the gold electrode surface via a reaction performed in the presence of ethyl(dimethylaminopropyl)carbodiimide / N-hydroxysuccinimide (EDC/NHS) between the amine group present in the probe and carboxylic groups present on the surface of the redox-active layer. Each modification step has been controlled with Osteryoung square-wave voltammetry. The genosensor incorporating the (dipyrromethene)2Cu(II) complex was able to detect a fully complementary single-stranded DNA target with a detection limit of 1.39 pM. A linear dynamic range was observed from 1 to 10 pM. This genosensor displays good discrimination between three single-stranded DNA targets studied: fully complementary, partially complementary (with only six complementary bases), and totally noncomplementary to the probe. When the (dipyrromethene)2Co(II) complex was applied, a detection limit of 1.28 pM for the fully complementary target was obtained. However, this genosensor was not able to discriminate partially complementary and totally noncomplementary oligonucleotide sequences to the probe. Electrochemical measurements, using both types of genosensors in the presence of different supporting electrolytes, were performed in order to elaborate a new mechanism of analytical signal generation based on an ion barrier “switch-off” system.

I

the detection of DNA by direct oxidation or reduction of nucleic acids using various electrodes have been published.7−14 In this type of sensor, the analytical signal results directly from redox reactions of some nucleotides; therefore, many authors focused their attention on the improvement of efficiency. The main drawback of this way of signaling is the background current at the relatively high potentials required for direct oxidation of DNA. One method that avoids this limitation relies on application of the same electrochemical mediators that facilitate electron transfer between the electroactive base and the electrode surface. This method permits sensitivity at the attomolar level in terms of concentrations of the target DNA.15 Genosensors based on free diffusional redox markers belong to analytical devices consisting of an electrode modified with label-free DNA probe and a redox marker in the sample

n the past decade, an impressive increase in the number of papers concerning the development of genosensors for DNA detection was observed.1−3 The common theme for all types of genosensors is a single-stranded DNA (ssDNA) attached to the surface of the transducer as a recognition element. The main differences between the genosensors rely on the signaling mechanism. High sensitivity, user-friendliness, relatively low price of a single analysis, and possibilities for miniaturization make the electrochemical genosensors very attractive tools suitable for medical diagnosis, environmental monitoring, and food quality control. Seminal work concerning the genosensors based on the direct electroactivity of nucleic bases was published by Paleček4 in a report concerning the electroactivity of DNA. The first electrochemical DNA sensor was developed by Millian and Mikkelson in 1993.5 Five years later, Paleček6 published a paper about the application of adsorption striping voltammetry for the determination of nanograms of DNA. From these first publications, during the next decades, many articles concerning © 2015 American Chemical Society

Received: May 13, 2015 Accepted: September 11, 2015 Published: September 11, 2015 9702

DOI: 10.1021/acs.analchem.5b01988 Anal. Chem. 2015, 87, 9702−9709

Article

Analytical Chemistry

approach does not work properly in complex samples. A sensor based on pseudoknot architecture, working according to a signal-on mechanism, was a good example of the improvement in analytical parameters.42,43 This E-sensor was able to detect the target in blood plasma with a detection limit of 2 nM and a useful dynamic range from 2 to 100 nM. Joining this research, we have developed a novel dual E-DNA sensor that generated the analytical signal according to a signaloff and signal-on scheme for simultaneous detection of two different sequences of DNA derived from avian influenza virus (AIV) type H5N1 by means of one electrode.44 Detection limits of this sensor were 40 nM and 20 nM for simultaneous analysis of both sequences, hemaglutinin and neuraminidase, respectively. These values are very promising from the diagnostic point of view. Recently in our laboratory, the click reaction between the amino group from the NH2-ssDNA probe and epoxy groups from AET/Phen-Epoxy/Fe(III) (Phen-Epoxy)2 complex immobilized on a gold electrode surface has been successfully applied for construction of the genosensor.45 The main advantages of this genosensor are its suitability for RNA determination and its ability to distinguish the different positions of the complementary parts present in RNA transcripts. In this paper, we have introduced a new type of sensitive l a b e l - f r e e e l e c t r o c h e m ic a l g e n o s e n s o r b a s e d o n (dipyrromethene) 2 Cu(II) and (dipyrromethene) 2 Co(II) redox-active layers. Their crucial feature is the location of the redox centers very close to the electrode surface. The ssDNA probes are located in line with the redox-active centers (Scheme S1). The mechanism of analytical signal generation by this new type of genosensor will be proposed and discussed.

solution. In this type of genosensor, the mechanism of analytical signal generation is connected with the binding event between the target and the probe located at the electrode surface. This process controls the access of the redox-active marker to the electrode surface. Finally, electron transfer from the marker to the surface of the electrode is affected by the hybridization of probe and target DNA strain. [Fe(CN)6]3‑/4‑ and [Ru(NH3)6]3+ complexes are most often used as the redoxactive marker ions.16−21 The detection of some mutated or damaged DNA bases is very important from the point of view of early diagnosis of genetic diseases. The genosensors have been successfully applied to study the genotoxicity and teratogenicity of small molecules such as pesticides.22−24 Application of small DNAintercalating or groove-binding electroactive compounds gives the possibility to distinguish a single-stranded probe from a double-stranded hybrid located at the surface of the electrode.25,26 One of the most commonly used intercalators successfully applied as a redox indicator of DNA duplex formation in genosensors is daunomycin.27−34 The main weak point of these two described types of sensors is the necessity of adding redox marker molecules to the sample solution. In 2003, Fan et al.35 described a new class of electrochemical DNA sensors (E-DNA). They used a single stem−loop DNA structure immobilized at the surface of a gold electrode via a S− Au bond. The unbound end was decorated with the redoxactive label ferrocene. Before hybridization, the stem−loop structure holds the ferrocene close to the electrode surface. After hybridization, the stem−loop is transferred to a linear duplex structure, and as a consequence, the distance between the electrode surface and the redox marker increases, leading to a decrease of the faradic current signal. The selectivity and sensitivity of E-DNA sensors arises from a combination of conformational change upon hybridization, together with the redox labels being active at potentials far from those of most electroactive biomolecules typical for clinical and environmental samples, thus being resistant to interfering contaminants. Generally, the genosensors based on this described mechanism of signaling belong to a very wide “signal-off” family.36−38 A limitation of signal-off sensors is that recognition processes are signaled by loss of the initial current. Thus, it is not possible to suppress the original current more than 100%. The first genosensor working according to a “signalon” mechanism was described by Immoos and co-workers.39,40 They described a sensor consisting of two strands of ssDNA. The capture and probe strand are linked together via a flexible linker of poly(ethylene glycol). The capture strand contained a 3′-terminal thiol for immobilization on the gold surface. The probe strand was decorated with ferrocene at the 5′-terminus. Hybridization with target DNA that was complementary to both immobilized strands moved the labeled end of the probe into close proximity to the electrode, increasing the electron transfer. This approach resulted in a detection limit of 200 pM and 600% current signal increase. Both type of E-DNA sensors, signal-on and signal-off, generated the analytical signal after stimulation with target DNA without the addition of exogenous reagents. They could be applied for measurements even in blood serum.41 This property is very important from the medical diagnostic point of view. A weak point of these types of E-DNA sensors is their rather complicated and not very stable architectures. This leads to poor reusability. Most signal-on systems contained signalgenerating strands noncovalently attached to the surface. This

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EXPERIMENTAL SECTION Preparation of Genosensor: Successive Steps of Gold Electrode Modification. Gold disk electrodes with a radius of 2 mm [Bioanalytical Systems (BASi), West Lafayette, IN] were used for the experiments. The electrodes were polished with wet 0.3 and 0.05 μm alumina slurry (α and γ micropolish; Buehler, Lake Bluff, IL) on a flat pad for at least 10 min and rinsed repeatedly with water. The polished electrodes were then dipped in a 0.5 M KOH solution deoxygenated by purging with argon for 15 min, and the potential was cycled between −400 and −1200 mV (versus Ag/AgCl reference electrode) with a scan rate of 100 mV·s−1; the number of cycles was 30, 50, or 10. Next, the electrodes were cleaned in 0.5 M H2SO4 solution in the potential window between 0.3 V and +1.5 V (versus Ag/ AgCl reference electrode) with a scan rate of 100 mV·s−1; the number of cycles was 3, 10, or 3. Before modification, the surfaces of electrodes was refreshed in 0.5 M KOH solution for 10 cycles. After the electrochemical cleaning was finished, each electrode was washed with Milli-Q water and stored in water (for several minutes, until the next step) to avoid contamination from air. Directly after cleaning, the electrodes were dipped into a dichloromethane modification solution based on a mixed solution of dipyrromethene (DPM) and 4-mercaptobutanol (MBT) in the molar ratio 1:100 (1 mM/100 mM) at room temperature for 3 h. Next, after being washed with dichloromethene, the electrodes were dipped in 1 mM Cu(II) or Co(II) acetate solution in chloroform and methanol (ratio 1:1 v/v) at room temperature for 1 h. Subsequently, the electrodes were dipped into a solution of 1 mM DPM-COOH 9703


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Figure 1. Representative Osteryoung square-wave voltammograms obtained for the gold electrode modified with (A) MBT/DMP-Cu(II)-DPMCONH-ssDNA and (B) MBT/DMP-Co(II)-DPM-CONH-ssDNA after hybridization with 20-mer complementary (c-NC3), partially complementary (pc-NC3), and noncomplementary (nc-NC3) ssDNA at concentrations (a) 0, (b) 1, (c) 2, (d) 4, (e) 6, (f) 8, and (g) 10 pM. Buffer conditions: 0.9 M NaCl and 0.09 M sodium citrate, pH 7.0.

step, these complexes were treated with DPM-COOH to saturate the coordination number of the metal centers. At the end of the modification procedure, the ssDNA probe was attached to the complex via an amide bond. Each of these steps was confirmed by the Osteryoung square-wave voltammetry (OSWV) method (Figure S1). To confirm that the complexes created are located at the surface of the electrode, the relationship between scan rate and current values has been tested (Figure S2). As can be seen, for the electrode modified with Cu(II) as well as Co(II) complexes, these relationships are linear for the anodic and cathodic current. This confirms the localization of appropriate redox centers on the surface of the electrode.46,47 The density Γ of redox-active layers DPM-Cu(II)-DPMCONH-ssDNA [(2.0 ± 0.07) × 10−11 mol·cm−2] and DPMCo(II)-DPM-CONH-ssDNA [(3.0 ± 0.06) × 10−11 mol·cm−2] assembled on the Au electrode surface were calculated on the basis of integration of the voltammetric peaks recorded at 100 mV·s−1 (Figure S2). The Γ value for DPM-Co(II)-DPMCONH-ssDNA was higher than that for DPM-Cu(II)-DPMCONH-ssDNA. These values are in a similar range as reported for other redox-active probes.48 The electrodes prepared in the above-described manner were applied for detection of totally complementary (c-NC3), partly complementary (pc-NC3), and noncomplementary (nc-NC3) ssDNA strains to the ssDNA probe, derived from avian influenza virus (AIV) type H5N1 (Figure S3). Figure 1 illustrates the representative OSWV spectra after incubation with the particular target ssDNA. The addition of target ssDNA caused a decrease in the faradic current. Noncomplementary ssDNA induced much smaller changes in the current (Figure 1). This phenomenon was observed for both examined genosensors. Our main goal was to develop the gensosensor with the best sensitivity, which might be used in the future as an alternative

in chloroform and methanol (ratio 1:1 v/v) for 18 h. In the next step, the carboxyl groups of DPM-COOH were activated by treatment with a mixture of 100 mM ethyl(dimethylaminopropyl)carbodiimide (EDC) and 50 mM Nhydroxysuccinimide (NHS) in 50 mM 2-(N-morpholino)ethanesulfonic acid (MES, pH 5.5) for 1 h to create the amide bond with amine groups of the NH2-ssDNA probe. Then, 10 μL droplets of 10 μM NH2-ssDNA probe in MES buffer, pH 7.0, were placed on the surface of each electrode. The electrodes were covered with tubes and stored for 6 h at 4 °C. After immobilization, the NH2-ssDNA probe electrodes were rinsed off with MES (pH 7.0) buffer, followed by quenching the remaining EDC/NHS-activated sites with 1 mM ethanolamine in 50 mM MES (pH 7.0) for 30 min. Then the modified electrodes were rinsed profusely with 50 mM MES (pH 7.0) and 0.9 M NaCl and 0.09 M sodium citrate (pH 7.0), respectively, and stored overnight at 4 °C until used. Detailed experimental data concerning materials and chemicals, electrochemical measurements, hybridization of NH2-NC3 probe and target sequences, and electrochemical characterization of the electroactive monolayer are described in the Supporting Information.

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RESULTS AND DISCUSSION At the beginning of our work, connected with preparation of the proposed electrochemical reagentless genosensors based on an ion-barrier switch-off system for sequence-specific detection of DNA, we have tested the effectiveness of step-by-step electrode modification with dipyrromethene Cu(II) and Co(II) complexes. Scheme S1 depicts in graphical form the modification scheme. The first step relies on the modification of gold electrode surface with a mixture of 4-mercaptobutanol and thiol derivatives of dipyrromethene. In the next step, Cu(II) or Co(II) complexes with dipyrromethene were formed at the surface of the electrode. In the following modification 9704


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Figure 2. Relative intensity of redox current in relation to (A) log concentration of 20-mer c-NC3 and (B) concentration of 20-mer pc-NC3 and ncNC3. (C) Concentrations of 20-mers: (■) c-NC3, (⧫) pc-NC3 and (●) nc-NC3. Buffer conditions: 0.9 M NaCl and 0.09 M sodium citrate, pH 7.0.

better analytical parameters than one incorporating DPMCo(II)-DPM. Its sensitivity in the picomolar range is comparable to the sensitivity reported for different genosensors based on redox-active self-assembled monolayers (Table S1). The main advantage of the sensing system proposed is the lack of the necessity for ssDNA strands labeling with redox-active probes. Also, this system has universal character. Any NH2ssDNA probe could be attached to the DPM-Me(II)-DPMCOOH layer via EDC/NHS coupling. Detection of mismatched base pairs in DNA plays a crucial role in the diagnosis of diseases, especially for early-stage treatment.54−56 The genosensor based on DPM-Cu(II)-DPM could be successfully applied for this purpose. Mechanism of Analytical Signal Generation. In the majority of recently developed electrochemical DNA or peptide nucleic acid (PNA) sensors, the redox centers are located at the end of the ssDNA strand directed toward the solution phase. In such a situation, the hybridization process alters the distance of the redox label from the electrode surface.40,41,44 In the proposed system, the redox centers are bound to the surface of electrode; thus changes of their distance to the electrode surface are not possible. In this aspect, the proposed sensors generate an analytical signal because of changes in environment around the redox center occurring as a result of hybridization processes. Scheme 1 illustrates the general redox processes going on in the redox centers of the described genosensors. For the sensor

to PCR sensing devices. Therefore, the possible lowest concentration range has been tested from 1 to 10 pM. The strongest responses were observed in the presence of fully complementary DNA target, for both type of genosensors (Figure 2A,C). But the response generated by the system incorporating DPM-Cu(II)-DPM was superior (ca. 80% current decrease) to the one based on DPM-Co(II)-DPM (ca. 50% current decrease). Weaker responses were observed in the presence of partially complementary target (pc-NC3), and the weakest responses were in the presence of noncomplementary sequence (nc-NC3) (Figure 2 B,C). Detection limits (DL) were calculated from49 DL =

3.3σ S

where σ is the standard deviation of the response and S is the slope of the calibration curve. DLs for the main c-NC3 target obtained in both systems were similar: 1.39 pM [recorded for DPM-Cu(II)-DPM] and 1.28 pM [recorded for DPM-Co(II)DPM]. Also, the linear dynamic range for the main target cNC3 obtained with both type of genosensors were very similar (Figure 2A,C). But the slope of genosensors incorporating Cu(II) sites was superior to the system based on DPM-Co(II)DPM. In addition, only this genosensor was able to discriminate pc-NC3 and nc-NC3 (Figure 2B). The signal-to-noise ratio (S/N) was estimated for genosensor incorporating DPM-Cu(II)-DPM, with the values 22.1, 16.0, and 4.6 obtained for c-, pc-, and nc-NC3, respectively. The S/N correlate well with strength of signals. In the presence of fully complementary DNA target, duplex was formed and its well-defined structure has the strongest influence on the redox center properties embedded into the sensing layer. Thus, we can say the value of S/N for fully complementary target is superior to that for totally noncomplementary one. The partially complementary target (pcNC3) has three complementary bases at the 5′-end and three complementary bases at the 3′-end (Figure S3). This is the reason for the high value of S/N and detection limit of 1.99 pM. The weaker interaction between NH2-NC3 ssDNA strand and ssDNA target having few complementary bases cannot be fully avoided.50−52 These phenomena also influence redox centers but to a smaller extent. Better selectivity might be obtained by use of electrolytes with different ionic strengths.53 We plan such work for validation of the proposed genosensors. The results presented in Figures 1 and 2 indicated that a genosensor based on DPM-Cu(II)-DPM complex displayed

Scheme 1. Schematic Illustration of Redox Reactionsa

a

(A) Occurring at gold electrodes modified with DPM-Cu(II)-DPM. (B) Occurring at gold electrodes modified with DPM-Co(II)-DPM complexes.

based on the Cu(II) complex, during the redox cycle the Cu(II) is reduced to Cu(I). As a consequence, a single minus charge of the reduced form appears at the surface of the electrode. The precondition of redox reaction run is the compensation of this charge by ions from the supporting electrolyte. In this case, cations will be involved. For the electrode modified with the Co(II) complex, an oxidation process is possible, which generates an extra positive charge of the oxidized form. For 9705


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Analytical Chemistry Scheme 2. Schematic Illustration of Signal Generation Mechanism of Genosensorsa

a

(A) Genosensors incorporating MBT/DPM-Cu(II)-DPM-CONH-dsDNA. (B) Genosensors incorporating MBT/DPM-Co(II)-DPM-CONHdsDNA.

Table 1. Oxidation and Reduction Peak Positions of Modified Gold Electrodesa

a Gold electrodes were modified with DPM-Cu(II)-DPM-CONH-ssDNA, DPM-Cu(II)-DPM-CONH-dsDNA, DPM-Co(II)-DPM-CONH-ssDNA, and DPM-Co(II)-DPM-CONH-dsDNA as indicated. Values were estimated in different electrolytes (1.0 M) + 0.01 M sodium citrate. Hydration energy (ΔhydrG) and radius of ions present in supporting electrolyte are also listed.68 Calculations were based on CV measurement at 100 mV/s.

its neutralization, anions from the supporting electrolyte will be involved. As a consequence, the transfer of electrons from the redox centers to the electrode surface and in the opposite direction is determined by the movements of ions from the supporting electrolyte at the border between the modified layer/electrolyte. Thus, it is worth noting that electron transfer from the redox centers to the electrode surface, which is essential for redox reactions, is determined by the availability of redox centers of appropriate ions. The important parameters that control this availability include the radius of hydrated ions,57−60 the structure of the redox center,61 and the structure of the border between the modifying layer and the electrolyte.62−64 The sensors under discussion consist of two basic elements: redox centers immobilized covalently to the surface of the electrode and ssDNA probes attached covalently to these redox centers (Scheme 2). In consequence, the redox centers are immersed inside a modifying layer and ssDNA probes are outside, directed toward the solution phase. The surface of electrode is negatively charged because of the presence of ssDNA. Of course, this charge is compensated by cations from the supporting electrolyte. After the hybridization reaction, the negative charge of the electrode surface is doubled, and also the effective thickness of the layer increases. The electric charge has a strong solvation effect, and it causes the

medium dielectric coefficient to decrease. Both factors lead to a lower capacitance for the hybridized surface.65−67 The changes of these parameters have an essential influence on the accessibility of ions to the redox centers.61,64 In consequence, this leads to changes of the kinetics of electron transfer from the redox centers to the electrode surface, and this is the basis of analytical signal generation by the described sensors. Taking into account the above hypothesis, we have checked the effect of electrolyte ions on the current values generated by the proposed genosensors. The anodic and cathodic potential peaks measured for electrodes modified with DPM-Cu(II)DPM-CONH-ssDNA, DPM-Cu(II)-DPM-CONH-dsDNA, DPM-Co(II)-DPM-CONH-ssDNA, and DPM-Co(II)-DPMCONH-dsDNA in the presence of different supporting electrolytes have been recorded. The results are collected in Table 1. For the electrode incorporating Cu(II), as a consequence of a reduction reaction, the complex became negatively monocharged (Scheme 1). This charge has to be compensated by a cation from the electrolyte. As we can see, the potential positions of the cathodic as well as anodic peak current were changed in relation to the applied electrolyte (Table 1). For the oxidation current peak, the potential values decrease according to the sequence Cs+ > K+ > Na+. This sequence follows the decrease of the cation hydration radius. In the case of 9706


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Analytical Chemistry Table 2. Electron Transfer Coefficients and Rate Constants Calculated for Modified Gold Electrodesa DPM-Cu(II)-DPM ssDNA −1

α CsCl KCl NaCl NaBF4 NaClO4

0.37 0.56 0.56 0.66 0.76

± ± ± ± ±

0.45 0.58 1.02 0.51 0.47

± ± ± ± ±

0.05 0.05 0.08 0.07 0.05

0.45 0.47 0.62 0.73 0.72

± ± ± ± ±

ssDNA −1

α

ks (s ) 0.04 0.01 0.08 0.04 0.04

DPM-Co(II)-DPM dsDNA α

ks (s ) 0.04 0.03 0.04 0.03 0.08

0.54 0.93 1.09 0.40 0.37

± ± ± ± ±

0.03 0.08 0.06 0.03 0.02

0.88 0.63 0.58 0.51 0.45

± ± ± ± ±

dsDNA −1

0.09 0.01 0.02 0.02 0.04

0.29 1.36 1.07 0.83 0.75

± ± ± ± ±

0.06 0.05 0.07 0.02 0.02

ks (s−1)

α

ks (s ) 0.62 0.64 0.55 0.60 0.40

± ± ± ± ±

0.05 0.03 0.02 0.03 0.06

0.93 0.88 1.15 0.82 0.81

± ± ± ± ±

0.07 0.04 0.06 0.03 0.05

a

Electrodes were modified with DPM-Cu(II)-DPM-CONH-ssDNA, DPM-Cu(II)-DPM-CONH-dsDNA, DPM-Co(II)-DPM-CONH-ssDNA, and DPM-Co(II)-DPM-CONH-dsDNA as indicated. Values were estimated in the presence of different supporting electrolytes (1.0 M) + 0.01 M sodium citrate.

The presented results indicate that in the processes of charge neutralization that arose at the border of the modified layer/ electrolyte solution during the redox reactions, not only the main cations or anions but also their counterions are involved. A similar phenomenon was described in the literature.57,69,70 In Table 2, values of the electron transfer coefficient (α) and electron rate constant (ks) calculated on the basis of the generally applicable Laviron’s procedure in the presence of different electrolytes are collected.71,72 For electrodes modified with DPM-Cu(II)-DPM, α values measured in the selected electrolytes increase in the order CsCl < KCl < NaCl < NaBF4 < NaClO4. A totally opposite tendency was observed for electrodes modified with DPM-Co(II)-DPM, where the α values decrease in the order CsCl > KCl > NaCl > NaBF4 > NaClO4. In most cases, the α values are >0.5, indicating that reduction is favored. Only in the case of the electrode modified with DPM-Cu(II)-DPM, with the most lipophilic cation Cs+, the calculated value is α < 0.5. This means that, in the presence of CsCl oxidation, Cu(I)/Cu(II) is favored. The opposite phenomenon was observed for the electrode modified with DPM-Co(II)-DPM. In the presence of CsCl the α value is >0.5, indicating a lower energy barrier for reduction. We have also tested the effect of anions on the values of α and ks for both types of electrodes studied. In the presence of ClO4−, the most lipophilic anions, the highest α values (0.76), indicating the lowest energy barrier for reduction, were obtained for the DPM-Cu(II)-DPM system. On the other hand, for the DPM-Co(II)-DPM system, the lowest energy barrier in the presence of ClO4− was for oxidation. In the presence of this electrolyte the lowest α values, 0.45, was obtained. The values of ks did not change in a regular manner. However, the highest ks value in the range of 1.0 s−1, was obtained in the presence of NaCl for both systems studied. Also, the highest value of current measured with OSVW was recorded in the presence of NaCl (Table S2). So these parameters confirmed the superiority of NaCl for the electrochemical measurements presented. Both systems studied have been explored in the presence of different electrolytes by use of Osteryoung square wave voltammetry, the main measuring technique applied for oligonucleotide sequence determination (Table S2). The potential values of the peak current recorded for DPMCu(II)-DPM increased in the order CsCl < KCl < NaCl < NaBF4 < NaClO4. On the other hand, for DPM-Co(II)-DPM, no regular changes of the potential values of the peak current were observed. For chloride salts, a decrease of this parameter was observed in the order CsCl > KCl > NaCl. For sodium salts

reduction, the opposite tendency was observed. The potential values of the peak current increase according to the sequence Cs+ < K+ < Na+. This tendency is the same for electrodes before and after hybridization. These observations justify that the reduction reaction Cu(II)/Cu(I) occurring at the surface of the electrode depends on neutralization of the negative charge by cations from the supporting electrolyte. The cations with a large radius and lower charge density penetrate the interior of the modified layer in which the redox centers are located and are able to efficiently neutralize their charge. This explains why the reduction reaction is running at a lower potential. On the other hand, the oxidation process leads to a neutralization of the immobilized complex and the release of cations from the modified layer to the bulk solution. This process is more difficult for more lipophilic cations. In the presence of the most lipophilic cation, Cs+, the oxidation of Cu(I)/Cu(II) was running at the highest potential (Table 1). In addition to different cations, Cs+, K+, and Na+ (chloride salts), we have tested different anions: Cl−, BF4−, and ClO4− (as sodium salts). The influence of different anions on the potential position of peak current is more significant in the case of the Cu(I)/Cu(II) oxidation process. The largest potential shift in positive direction was observed in the case of perchlorate anions. In the case of Cu(II)/Cu(I) reduction, the tendency was the same, but the potential shift was smaller. After hybridization, the same tendencies, albeit weaker, were observed for both processes (Table 1). For the electrode modified with the DPM-Co(II)-DPM complex, the first possible process is oxidation to Co(III), which generated a monocation complex (Scheme 1). Such charge has to be neutralized by anions from the supporting electrolyte. As was observed for the DPM-Cu(II)-DPM complex, the influence of the electrolyte ions was significant, but the potential shifts of the oxidation and reduction peak current were smaller. The observed potential shift tendency was as follows: Cs+ < K+ ≤ Na+ for oxidation, and Cs+ ≥ K+ > Na+ for reduction. After hybridization, these tendencies remained the same, but the peak potential shifts were lower for both processes. These tendencies concerning different cations observed for the electrode modified with DPM-Co(II)-DPM were in the opposite direction in comparison to those for the electrode modified with the DPM-Cu(II)-DPM complex. On the other hand, the potential shift of peak current for both oxidation and reduction processes recorded in the presence of different anions for all electrodes studied was the same: Cl− < BF4− < ClO4−. The lowest energy barrier for electron transfer between electrode−redox center occurred in the presence of the smallest and most hydrophilic Cl− anions (Table 1). 9707


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ssDNA probes for designing of genosensors, which is crucial from a medical diagnostic point of view.

with more lipophilic anions, NaBF4 and NaClO4, the potential of the peak current was similarly in the 380−400 mV range. Generally, it is visible that values of the current signal generated by the electrodes modified with redox-active layers are regulated by the composition of main cations or anions and their proper counterions from the supporting electrolyte and their affinity to the redox centers. The lower sensitivity and selectivity of the genosensor based on the DPM-Co(II)-DPM complex is a consequence of the relatively easy access of electrolyte ions to the Co(II) redox center. This means that, in the case of the DPM-Co(II)-DPM complex, the energy barrier of accessibility to the redox center for electrolyte ions is relatively low. As a consequence, in this case, in order to generate the analytical signal it is necessary to cause more significant changes around the redox centers. This means that it is necessary to bind to the surface of electrode a higher number of ssDNA strands in comparison to the electrode modified with DPM-Cu(II)-DPM complex, where the energy barrier of accessibility to the redox center for electrolyte ions is higher and even small obstacles in the way of the ions to the redox center cause signal generation. These differences in the accessibility of ions to the redox centers originate from the structure of the complexes. The DPM-Co(II)-DPM complex is planar,73,74 whereas the DPMCu(II)-DPM complex is twisted.74,75 Because of this, Cu(II) is located in a small hydrophobic cavity. This causes an increase in the energy barrier for access of the electrolyte ions. In consequence, even small numbers of bonded target ssDNA were able to generate a strong analytical signal. The location of the redox center in the hydrophobic cavity stabilized the DPMCu(I)-DPM complex, which in water solution is unstable.76 Thanks to that, our system is reversible.

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ASSOCIATED CONTENT

* Supporting Information S

Experimental data and additional figures, scheme and tables. The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.analchem.5b01988. Detailed experimental data concerning materials and chemicals, electrochemical measurements and characterization of the monolayer, and hybridization of probe and target sequences; a scheme showing steps of genosensor preparation; five figures showing representative Osteryoung square-wave voltammograms and cyclic voltammograms and targets of 20-mer ssDNA; two tables listing comparison of electrochemical genosensors and electrochemical parameters (PDF)

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AUTHOR INFORMATION

Corresponding Author

*E-mail [email protected]; tel +48 89 523 46 12; fax +48 89 524 01 24. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This work was supported by Innovative Economy Program WND-POIG.01.01.02-00-007/08 and 679/N-BELGIA/2010/ 0, by COST Action CM1005 “Supramolecular Chemistry in Water”, and by statutory fund of Institute of Animal Reproduction and Food Research of Polish Academy of Sciences, Olsztyn, Poland.

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CONCLUSIONS In this paper we have introduced a novel type of electrochemical label-free genosensor based on DPM-Cu(II)-DPMCOOH or DPM-Co(II)-DPM-COOH redox-active monolayers suitable for amide coupling of ssDNA strands. In this approach, the redox centers, responsible for analytical signal transduction, are located very close to the electrode surface and their positions remained unchanged during the entire analytical procedure. In spite of this, changes in redox activity of Cu(II) or Co(II) centers were observed upon hybridization. The mechanism of this unusual analytical signal generation relies on changes in accessibility of ions present in the supporting solution to the redox centers in order to neutralize the charge occurring as a result of oxidation/reduction processes. This phenomenon depends not only on the hydrophobicity/ hydrophilicity of ions but on the structure of redox complexes as well. Better analytical parameters have been achieved with DPM-Cu(II)-DPM than with DPM-Co(II)-DPM. The twisted structure of DPM-Cu(II)-DPM created a higher energy barrier for supporting electrolyte ions in comparison to the flat structure of DPM-Co(II)-DPM. A novel mechanism of electrochemical signal generation based on changes of the ion-barrier “switch off” system have been proposed. The proposed genosensors have been tested for determination of oligonucleotide sequences specific for avian influenza virus H5N1 with detection limit in the picomolar range. The applied DPM-Me(II)-DPM self-assembled monolayer functionalized with carboxylic groups has universal character and could be used for amide coupling of any NH2-

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