Chemical Control of Electrode Functionalization for Detection of DNA

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Langmuir 2005, 21, 1022-1027

Chemical Control of Electrode Functionalization for Detection of DNA Hybridization by Electrochemical Impedance Spectroscopy Shanlin Pan and Lewis Rothberg* Department of Chemistry, University of Rochester, Rochester, New York 14627 Received July 29, 2004. In Final Form: October 31, 2004 We report sensitive label-free detection of DNA oligonucleotide sequences using ac impedance measurements. The surface attachment chemistry is critical, and using mixed self-assembled monolayers on a gold electrode results in much better performance than homogeneous self-assembled monolayers. Contrary to expectations, binding of the target sequence reduces rather than increases the charge transfer resistance. Similar behavior is observed on indium tin oxide electrodes, and we ascribe it to the hydrophilicity and rigidity of the DNA duplex that cause it to reside further from the electrode surface and facilitate the approach of negatively charged redox moieties to the interface.

Introduction There is considerable interest in electrical sensing of biomolecular binding since it has the potential to be label free, to work easily in aqueous environments native to the biomolecules, and to be integrated with small, fast, and inexpensive microelectronics as detection instrumentation. Various electrochemical approaches to DNA detection have been developed1 for detection of oligonucleotide sequences2,3 and DNA damage.4 Electroactivity of the nucleic acids themselves,4-9 incorporation of electroactive markers2,3,10,11 onto the nucleic acids, label-free detection methods using redox reactions modified by DNA hybridization,12-15 and selective intercalation of electroactive moieties into duplex DNA16 have all been demonstrated. A variety of electrochemical measurement protocols have been used including cyclic voltammetry,5 stripping potentiometry,11 square wave voltammetry,4 differential voltammetry,7 and ac impedance spectroscopy.12-15 * Corresponding author. E-mail address: rochester.edu.

rothberg@chem.

(1) Palecek, E. Talanta 2002, 56, 809-819. (2) Choong, V.; Gallagher, S.; Gaskin, M.; Li, C.; Maracas, G.; Shi, S. US patent PCT, WO 01/42508. (3) Pividori, M. I.; Merkoci, A.; Alegret, S. Biosens. Bioelectron. 2000, 15, 291-303. (4) Mugweru, A.; Rusling, J. F. Anal. Chem. 2002, 74, 4044-4049. (5) De-los-Santos-Alvarez, P.; Lobo-Castaqsn, M. J.; Miranda-Ordieres, A. J.; Tuqsn-Blanco, P. Anal. Chem. 2002, 74, 3342-3347. (6) Sistare, M. F.; Holmber, R. C.; Thorp, H. H. J. Phys. Chem. B 1999, 103, 10718-10728. (7) Olivira-Brett, A. M.; Silva, L. A.; Brett, C. M. A. Langmuir 2002, 18, 2326-2330. (8) Armistead, P. M.; Thorp, H. H. Anal. Chem. 2000, 72, 37643770. (9) Thorp, H. H. Trends Biotechnol. 1998, 16, 117-121. (10) Yu, C. J.; Wan, Y. J.; Yowanto, H.; Li, J.; Tao, C.; James, M. D.; Tan, C. L.; Blackburn, G. F.; Meade T. J. J. Am. Chem. Soc. 2001, 123, 11155-11161. (11) Wang, J.; Xu, D. K.; Kawde, A. N.; Polsky, R. Anal. Chem. 2001, 73, 5576-5581. (12) Ruan, C. M.; Yang, L. J.; Li, Y. B. Anal. Chem. 2002, 74, 48144820. (13) Yan, F.; Sadik, O. A. Anal. Chem. 2001, 73, 5272-5280. (14) Patolsky, F.; Katz, E.; Bardea, A.; Willner, I. Langmuir 1999, 15, 3703-3706. (15) Patolsky, F.; Lichtenstein, A.; Willner, I. J. Am. Chem. Soc. 2001, 123, 5194-5205. (16) Drummond, T. G.; Hill, M. G.; Barton, J. K. J. Am. Chem. Soc. 2004, 126, 15010-15011.

In the present work, we use ac impedance measurements to detect changes at the electrode-solution interface caused by oligonucleotide hybridization. Our approach is inspired in large measure by the considerable success that ac impedance measurements have had in electrical characterization of self-assembled monolayers (SAMs).17-18 The method is simple, sensitive, requires no labeling of the analyte with a redox active moiety, and works well in the highly ionic environment appropriate to biomolecular detection. The focus of the present paper is to show how ac impedance measurements can be improved in sensitivity by careful control of the attachment chemistry at the working electrode surface. Optimization of the detection of target binding requires that the probe not pack too densely, that there be reasonable amounts of current flow through the attachment layer, and that this current is modulated by the binding event. We show that the use of mixed self-assembled monolayers for attachment of the probes is an effective strategy for satisfying these criteria. The probe density is regulated by controlling the relative amounts of chemically inert and active linkers in the monolayer. When the active linkers create imperfections in the SAM or serve to block charge transfer in an already leaky SAM, target binding to the active linkers can be effective in producing charge-transfer impedance changes that form the basis for sensing. In the present paper, we use mixed SAMs of 2-mercaptoethanol and 11-mercaptoundecanoic acid on a gold working electrode. Probe oligonucleotides are bound to the activated undecanoic acid through a robust streptavidin-biotin chemistry. ac impedance measurements are made using [Fe(CN)6]4-/3- redox couple as a signal reporter for surface bonding events to enhance the sensitivity of the method. We can reproducibly detect binding of 10 pmol of target sequence while noncomplementary sequences do not affect the measured impedance. Since the redox currents we work with are quite large, subfemtomolar sensitivity should be achievable simply by scaling to smaller electrodes. Generally, adding material tends to obstruct charge flow, but we find that target binding (17) Janek, R. P.; Fawcett, W. R.; Ulman, A. Langmuir 1998, 14, 3011-3018. (18) Nicole, M. J.; Michael, G. H. Curr. Opin. Chem. Biol. 2001, 5, 209-215.

10.1021/la048083a CCC: $30.25 © 2005 American Chemical Society Published on Web 01/06/2005

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Figure 1. Scheme for the surface functionalization of the gold electrode for DNA sequence detection. Step I: SAM formed from mixed thiols 11-MUA and 2-ME. Step II: Streptavidin attachment to 11-MUA carboxylate group. Step III: Biotinylated oligonucleotide probe sequence attachment (biotin-5′-TTT TTT TTT GTT CTT CTC ATC ATC-3′). Step IV: Target sequence (5′-GTT CTT CTC ATC ATC-3′) binding. Detailed experimental conditions can be found in the text.

actually decreases the charge-transfer resistance and we speculate on the reasons. Experimental Procedures Gold Electrode Preparation. Glass substrates (1.8 cm × 1.8 cm) were cleaned in piranha solution (v/v 3:1 H2SO4/H2O2) at 100 °C for 1 h, rinsed thoroughly in deionized water, and dried under flowing nitrogen gas. Gold electrodes were prepared by evaporation of 5 nm of Cr followed by 200 nm of Au in a thermal evaporator with base pressure ∼10-6 Torr. Four electrodes of 0.25 cm2 each and connections to the edge of the substrate for each of these were deposited simultaneously using a shadow mask. The substrates were then cleaned by sonication in highly purified ethanol for 5 min and dried under flowing nitrogen gas before use. The connections to the edge were covered with an insulating layer formed by half epoxy resin and half ethylenediamine cured for 3-5 h. The resin serves both to define the electrode size precisely and to localize the analyte solution on the contacts. Curing was followed by another rinse with deionized water and drying under flowing nitrogen gas. Surface Functionalization of Gold Electrodes. The attachment chemistry used on the electrodes is illustrated schematically in Figure 1. Self-assembled monolayers were formed by immersion of the electrodes for 24 h in methanol solutions of 2-mercaptoethanol (2-ME), 11-mercaptoundecanoic acid (11MUA), or mixtures of the two of these. The ratios were varied to optimize the sensor performance as will be discussed in the results section. For the remaining chemistry, aqueous solutions with 0.5 M NaCl (pH ) 7.5 with 10 mM phosphate buffer) that we refer to as saline solution were used. Attachment of streptavidin was accomplished by immersion of the SAM-coated electrodes in saline solution containing 50 mg/mL 1-ethyl-3-(3-diamethylamine-propyl) carbodiimide for 12 h followed by immersion for 7 h in saline solution with 0.05 mg/ mL streptavidin. The substrate is then rinsed with saline solution. Commercially obtained biotin-modified oligonucleotide probe strands (biotin-5′-TTT TTT TTT GTT CTT CTC ATC ATC-3′, MWG Biotech) were grafted onto the streptavidin by applying 0.1 mL drops of saline solution with 10 µM probe concentration for 5 h. Finally, an analyte consisting of unmodified DNA oligonucleotides in saline solution was applied for 7 h in a humid environment to prevent evaporation while hybridization was being attempted. Either a complementary single-stranded DNA target (5′-GAT GAT GAG AAG AAC-3′, 10 µM) or a noncomplementary control (5′-GTT CTT CTC ATC ATC-3′, 10 µM) sequence was used as the analyte. Indium Tin Oxide Electrode Preparation and Functionalization. Commercial substrates with indium tin oxide (ITO) films on glass (30-60 Ω/square resistivity) were cut into 0.5 cm by 2 cm slides and cleaned by sequential rinsing in acetone, ethanol, and water. The slides were immersed in 1 M HCl for 10 min and then treated by 1:1:5 (v/v) H2O2/NH4OH/H2O for 1 h to fully hydroxylate the surface. The substrates were thoroughly rinsed with deionized water and dried under flowing nitrogen gas. Their surfaces were modified using a 1% solution of (3-

glycidyloxypropyl) trimethoxysilane in toluene for 24 h followed by rinsing with toluene and drying under flowing nitrogen gas. Half of the ITO electrode was coated with epoxy resin as with the gold and used to make connections, leaving the remaining half (0.5 cm × 1 cm) to be used as the working electrode. The working electrodes then were treated directly by 0.05 mg/mL streptavidin, and the remaining procedures were identical to those used with the gold electrodes. ac Impedance Measurements. Following each stage of the chemical functionalization, ac impedance spectra and cyclic voltammograms were taken using a BAS 100 B/W (Bioanalytical Instruments) electrochemical setup with a platinum wire as counter electrode and a standard Ag/AgCl reference electrode. The counter electrode and reference electrode were bound together and put directly on top of the functionalized evaporated electrodes. The electrolyte was 0.5 M buffered NaCl solution at room temperature containing 10 mM potassium hexacyanoferrate(II/III), K4Fe(CN)6/K3Fe(CN)6. Cyclic voltammograms were taken in the same electrolyte solution. All impedance spectra were taken using ac modulation of 5 mV over the frequency range from 0.1 Hz to 1 kHz with 280 mV dc bias with respect to a Ag/AgCl reference electrode. The choice of 280 mV potential is near equilibrium where reductive and oxidative rates for the [Fe(CN)6]3-/[Fe(CN)6]4- pair are equal so that the redox species will not be depleted near the electrode surface during the ac impedance measurements. Using the relationship between the peak current of the cyclic voltammogram Ip, the scan speed ν, the number of electrons exchanged n, and the known diffusion coefficient DR and concentration CR of the reduced species, we calculate the effective area of the bare gold electrodes to be 0.30 cm2, in good agreement with the fabricated geometry.18

Results and Discussion Theory and Interpretation of ac Impedance. The information of interest about the surface chemistry is derived by modeling the charge flow to the electrode with an equivalent linear circuit as in the work of Randles.19,20 We use the equivalent circuit shown in Figure 2A as an approximate model for our ac impedance data. The interfacial layer at the working electrode is represented by the parallel combination of resistance RCT, which models charge transfer across the functionalized layer, and capacitance CDL, which accounts for the electrical double layer that forms at the interface. The series impedance ZW is the standard Warburg term that describes concentration depletion of the redox species near the interface.18,20 (19) Steel, A. B.; Levicky, R. L.; Herne, T. M.; Tarlov, M. J. Biophys. J. 2000, 79, 975-981. (20) Bard, A. J.; Faulkner, L. R. Electrochemical methods: fundamentals and applications, 2nd ed.; John Wiley and Sons: Hoboken, NJ, 2001.

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Figure 2. (A) The Randles circuit model for a modified gold electrode in an electrolyte. WE, CE, and RE denote working electrode, counterelectrode, and reference electrode, respectively. RCT and RS are interfacial charge transport resistance and uncompensated solution resistance, respectively, while CDL is capacitance associated with the charge double layer at the modified electrode. ZW is the Warburg impedance describing depletion of the redox species in the interfacial region as described in the text. (B) Typical Nyquist plot of reactive (Z′′) versus resistive (Z′) part of the complex impedance Z(ω) ) Z′ + jZ′′ (where j ) (-1)1/2) expected for a reversible electrochemical reaction that is diffusion controlled at low ac modulation frequency ω and electron-transfer controlled at high modulation frequency ω. This figure is adapted from those in ref 20.

The circuit model of Figure 2A leads to equivalent impedance of

Z(ω) ) Z′ + jZ′′

(1)

where

Z′ ) Rs +

RCT + σω-1/2 (CDLσω1/2 + 1)2 + ω2CDL2(RCT + σω-1/2)2 (2)

and

Z′′ ) Rs + ωCDL(RCT + σω-1/2)2 + σω-1/2(CDLσω1/2 + 1) (CDLσω1/2 + 1)2 + ω2CDL2(RCT + σω-1/2)2

(3)

In the above expressions,20 ω is the angular frequency of the ac modulation, RCT ) RT/(nFIo) where R is the gas constant, T is temperature, n is the number of moles of electrons exchanged in the redox reaction, F is Faraday’s constant, and Io is the exchange current density, directly related to the rate constant k0 for electron transfer. Rs is the uncompensated solution resistance as indicated in Randles circuit model of Figure 2. The Warburg impedance, ZW, and σ are related as follows

ZW ) σω-1/2(1 - j) and

σ)

(

)

1 1 RT + 2 1/2 1/2 1/2 n F A2 DO CO* DR CR* 2

(4)

where j ) (-1)1/2, A is electrode area, Di are species diffusion coefficients, and Ci* are bulk concentrations of redox species where O refers to the oxidized and R to the reduced species. ac impedance spectra measure Z′ and Z′′ versus ω and can be plotted as Nyquist diagrams (Figure 2B) that typically exhibit two important regions for a reversible reaction at solid interface. As indicated in the sketch, the semicircle obtained at high modulation frequency describes the faradic electron-transfer process at the electrode while the straight region obtained for lower modulation frequency contains information about the diffusionlimited transport of the redox species in the electrolyte to the electrode interface. It is straightforward to fit data to eqs 2 and 3 to extract the critical electrical properties of the interface, RCT and CDL. These parameters are very

Figure 3. Nyquist plots for mixed monolayers assembled on gold from solutions with different ratios of 2-mercaptoethanol (2-ME) and 11-mercaptoundecanoic acid (11-MUA). In each case, the total concentration of the solutions is 1 mM. Impedance spectra for the bare gold electrode (b), pure 11-MUA layer (2), and mixed monolayers with 1:1 (1) and 1:5 (O) 11-MUA:2-ME are shown.

sensitive to the biomolecular binding events at the modified electrode that we wish to detect. The solution resistance between reference electrode and working electrode, RS, and the Warburg impedance can also be obtained. Electrode Impedance Changes for Mixed SAMs on Gold. 11-MUA molecules can form more dense and ordered monolayers on gold that exhibit higher impedance than 2-ME as illustrated in Figure 3. The impedance is dominated by slow long-range charge transfer across the SAM with less defect, and the charge-transfer resistance RCT is larger than 1 MΩ. We require small RCT so that further chemical modification of the interface appropriate to sensing substantially affects the electron-transfer rate at the electrode. We can achieve this by forming the SAM with mixed thiol derivatives and Figure 3 shows the results of using 2-ME along with 11-MUA at 1:1 and 5:1 by concentration. The mixed monolayers have reduced impedance and higher current with the 5:1 samples exhibiting textbook behavior so that RCT, CDL, ZW, and RS are easily and uniquely extracted from the data. The dilution of the 11-MUA whose carboxylate end groups are used for biomolecular attachment is also a benefit of this strategy as it reduces the steric problems associated with hybridization (or other sensing chemistry) in a constrained environment.

Control of Electrode Functionalization

ac impedance spectra of pure 2-ME SAMs on gold are similar to those of the bare gold electrode. This indicates either that there are large numbers of defects in those SAMs so that leakage through pinholes in the SAM allows close approach of the redox agent to the bare gold or, more probably, that charge transfer across 2-ME is relatively facile. The electronic tunneling coefficient (β) is around 1 per methylene,21-22 permitting sufficiently rapid electron tunneling through 2-ME so that it is not rate limiting for the [Fe(CN)6]4-/[Fe(CN)6]3- redox chemistry. Under these circumstances, one might conjecture that the current would not be modulated by attachment chemistry on the 11-MUA sites. As shown below, however, we observe significant modulation of the electrode resistance by further chemical attachment. Either the pinholes in the primarily 2-ME layer are well correlated to 11MUA binding sites or, more likely, 20% coverage of the much larger 11-MUA is adequate to impede the approach of the redox compound to the electrode surface. Thus, modifications at the carboxylic acid terminus of the 11MUA are effective in further blocking the flow of the [Fe(CN)6]4-/3- redox moiety to the interface and therefore effective in modulating the current. It is worth noting that analogous ac impedance measurements without the [Fe(CN)6]4-/3- that use the NaCl in solution as an electrolyte resemble the pure 11-MUA and do not change significantly upon further chemical attachment to the 11MUA. We infer that thiol loss during the process of taking the impedance spectra is negligible since the cyclic voltammograms are very stable even after tens of cycles. When we deliberately strip the secondary binder layers and electrochemically desorb thiol from the gold electrode in concentrated KOH solution under negative potential, the desorption current is the same before and after the series of treatments and impedance measurements mentioned above. Linker and Probe Attachment. Nyquist plots after the various stages of chemical functionalization depicted in Figure 1 are shown in Figure 4A along with fits to the theory to extract circuit element parameters for the Randles model. Table 1 summarizes the results of fitting the impedance spectra corresponding to the different binding steps to the Randles model. Each step in the chemistry leads to increases in RCT as would be expected since additional bulk at the interface will impede charge transfer between the electrode and the redox agent. This, of course, bodes well for using ac impedance to sense biomolecular attachments at analogous surfaces. The solution resistance RS, dominant in the limit of high ac modulation frequency (cf. Figure 2B), is reasonably constant. This is also as expected since we are only modifying the interfacial layers. The jump of RS from 59 to 170 Ω for the probe modification step is due to changes between reference and working electrode since their distance is controlled manually. Cyclic voltammetry (CV) data versus sweep rate ν for the electrode after DNA probe attachment (step III) are illustrated in Figure 4B. The positive and negative peaks are due to the reduction and oxidation reaction of the [Fe(CN)6]4-/3- redox pair, respectively. The peak-peak separation and peak broadening with sweep rate derive from irreversible electrochemical reactions that occur at higher sweep rate. The peak currents ip for both oxidation and reduction follow a linear relation with v0.5 (inset) exhibiting diffusion-controlled electrochemical reaction (21) Xu, J.; Li, H. L.; Zhang, Y. J. Phys. Chem. 1993, 97, 1149711500. (22) Zamborini, F. P.; Leopold, M. C.; Hicks, J. F.; Kulesza, P. J.; Malik, M. A.; Murray, R. W. J. Am. Chem. Soc. 2002, 124, 8958-8964.

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Figure 4. (A) Nyquist plots of the sensing electrode response at different stages in the electrode functionalization process shown in Figure 1. Bare gold electrode (b), after 1:5 11-MUA: 2-ME mixed monolayer (O), after streptavidin attachment (4), and after biotinylated probe oligonucleotide attachment (0). All spectra are recorded in the presence of 10 mM (Fe(CN)6)3-/4in 0.5 NaCl as a redox-active indicator. Fits to the Randles model (eqs 2 and 3) are the solid lines, and the best fit parameters are reported in Table 1. (B) Cyclic voltammograms of DNA probe sequence functionalized gold electrode after step III taken at different voltage sweep rates (25, 50, 100, 250, and 500 mV/ s). The inset illustrates the peak currents scaling as ν1/2 (ν is potential sweeping rate in mV/s) reflecting current limited by diffusion of the redox species to the electrode.

rates for the redox pair at the electrode. The diffusioncontrolled behavior of the cyclic voltammogram indicates that there is negligible adsorption of redox species on the functionalized gold electrode surface and the redox reaction is close to diffusion-controlled reversible reaction since the peak to peak separation is close to 60 mV for slow sweep rate. Detection of Oligonucleotide Targets on Gold. The probe-functionalized surfaces are exposed to either complementary or noncomplementary DNA oligonucleotides under conditions suitable for hybridization. The resulting electrical responses are depicted in Figure 5 with fits to the Randles model. The corresponding fit parameters are listed in Table 1 and very little change in the model circuit parameters is observed upon exposure to the noncomplementary sequence while hybridization with the complement has a strong effect, primarily decreasing RCT and increasing CDL. The decrease in charge-transfer resistance RCT with hybridization is surprising in light of the monotonic increase in resistance observed in the other functionalization steps.

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Table 1. Best Parameters for Fitting Impedance Spectra Measured in Figures 4, 5, and 6 to the Randles Model (eqs 2 and 3)a sample modification bare gold step I step II step III step IV C-target step IV NC-target bare ITO probe/ITO C-target ITO a

CDL (F) 10-5

10-5)

6× (1 × 7 × 10-5 (1 × 10-5) 7 × 10-5 (1 × 10-5) 8.7 × 10-6 (0.6 × 10-6) 2 × 10-5 (2 × 10-5) 1 × 10-5 (2 × 10-6) 9 × 10-5 (1 × 10-5) 4 × 10-5 (5 × 10-6) 5 × 10-5 (1 × 10-5)

RCT (Ω)

σ (Ω Hz1/2)

RS (Ω)

40 (5) 116 (5) 295 (5) 440 (10) 80 (25) 430 (10) 20 (1) 72 (2) 28 (3)

76 (1) 89 (2) 90 (2) 780 (30) 510 (10) 660 (10) 63 (1) 113 (2) 126 (2)

69 (5) 58 (3) 59 (3) 170 (10) 180 (40) 165 (30) 51 (2) 59 (2) 57 (4)

Values in parentheses are estimates of the fit parameter accuracy.

Figure 5. ac impedance response before (0) and after (9) the probe sequence functionalized electrode is treated by (A) complementary (5′-GAT GAT GAG AAG AAC-3′) target sequence and (B) noncomplementary sequence 5′-GTT CTT CTC ATC ATC-3′. The solid lines are fits to the Randles model with best parameters listed in Table 1. Target concentrations are 10 µM, and the total amount of target in the applied analyte is 10 pmol.

Given our hypothesis that the primary current under our conditions comes from electron transfer to the redox agent at 2-ME domains or pinholes in the attachment SAM, this result suggests that hybridization acts to expose rather than obscure these locations. One possible reason is that double-stranded DNA is both stiffer and more hydrophilic than single-stranded DNA. This could serve to make the DNA reside more orthogonal to and further from the electrode surface,8 geometric changes that would facilitate approach of the redox species to the electrode. This speculation is supported by the large changes in σ (i.e., ZW) that reflect less redox species depletion in the interfacial region in the presence of the hybridized double-

stranded DNA. When the redox species more easily penetrates the interfacial layer (low σ and ZW), it can find places closer to the electrode for the electron transfer and may underlie the observed reduction in RCT. We do not fully understand the increase in double layer capacitance upon hybridization, but it is plausible that increased phosphate backbone charge will accommodate the accumulation of more counterions and a denser double layer near the electrode interface. It is tempting to associate a decrease in charge-transfer resistance upon duplex formation with previous reports23-28 that postulate improved charge transport in doublestranded relative to single-stranded DNA. However, Figure 3 shows that 11-MUA domains where DNA is attached contribute negligible redox current. Hence, changes in conductivity through the DNA itself are almost certainly irrelevant since almost all of our redox reactions occur across the very thin 2-ME domains where Fe(CN)64-/3can approach the gold. In the present case, we have no evidence for selective intercalation of the redox complex into the duplex and it would be very surprising if this happened. We have deliberately tried to saturate the binding sites by applying 10 pmol of target oligonucleotide, approximately an order of magnitude more than can be accommodated by the surface. Assuming 20% coverage of 11MUA, the measured 0.3 cm2 electrode area should have its binding sites saturated by approximately 1 pmol of oligonucleotide. This represents the worst case sensitivity of the detection and we think it is conservative based in Figure 5 to estimate that we can observe resistance changes an order of magnitude smaller than saturation. Since we are working with microampere currents at this electrode size, we also believe that it is conservative to estimate that we can shrink our electrode area by 3 orders of magnitude (to ∼150 µm × 150 µm). Thus, straightforward scaling should enable us to detect 100 amol (10-16 mol) of oligonucleotide target with this quite simple arrangement. Detection of Oligonucleotide Targets on ITO. Figure 6 presents data analogous to those of Figures 4 and 5 where we use an ITO electrode and silanization chemistry to attach DNA probes. In this case, the surface attachment SAM is sufficiently leaky so that the mixed monolayer strategy is not necessary to obtain substantial currents. We are also able to detect changes in the electrical (23) Boon, E. M.; Barton, J. K. Bioconjugate Chem. 2003, 14, 11401147. (24) Boon, E. M.; Jackson, N. M.; Wightman, M. D.; Kelley, S. O.; Hill, M. G.; Barton, J. K. J. Phys. Chem. B 2003, 107, 11805-11812. (25) Boon, E. M.; Barton, J. K. Curr. Opin. Struct. Biol. 2002, 12, 320-329. (26) Porath, D.; Cuniberti, G.; Di Felice, R. Top. Curr. Chem. 2004, 237, 183-227. (27) Conwell, E. M. Top. Curr. Chem. 2004, 237, 73-102. (28) Jackson, N. M.; Hill, M. G. Curr. Opin. Chem. Biol. 2001, 5, 209-215.

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Conclusions and Prospects Interfacial morphology is critical for making good sensing layers appropriate to electrochemical detection. We have demonstrated how mixed self-assembled monolayers on gold can be used to adjust the electrical properties to be sensitive to binding at the interface. In addition, the mixed monolayer strategy allows for dilution of probes to reduce steric crowding that makes target binding in restricted geometries slow and inefficient. Using ac impedance measurements and a redox agent, we have detected picomole quantities of DNA oligonucleotide binding at an electrode but reasonable scaling to smaller electrodes suggests it should be straightforward to detect 10-16 mol. The attachment chemistry for the probes is based on biotin streptavidin recognition and is therefore robust and generally applicable to large numbers of commercially available biotinylated proteins and antibodies. It should also be easy to make these measurements in an array format and to use mixed analytes for simultaneous parallel detection. Because the method is electrical, a sensor with integrated microfluidics and electronics could be engineered to be small and portable.29-30

Figure 6. (A) Nyquist plots of bare ITO sensing electrode (b) and the same electrode after DNA probe attachment (0) and exposure to a complementary target (9). Sequences, concentrations, and the electrolyte are the same as for Figures 4 and 5. The solid lines are fits to the Randles model (eqs 2 and 3) and the best fit parameters are reported in Table 1. (B) Cyclic voltammograms of DNA probe sequence functionalized ITO electrode taken at different voltage sweep rates (25, 50, 100, 250, and 500 mV/s). The inset illustrates the peak currents scaling as ν1/2 (ν is the potential scanning rate) reflecting current limited by diffusion of the redox species to the electrode.

properties of the interface upon target binding. Again, we have fit the ac impedance data with the Randles model in Table 1. Unfortunately, it is probably not as promising as the gold for use as a sensor. The functionalization chemistry is much less stable than that on gold and there is much less control of the probe density than with the mixed monolayer approach. The primary reason we include the ITO data is to show that the anomalous decrease in resistance RCT upon hybridization of the target DNA is also observed in this case. We assume that the same explanation we have proposed above also applies here.

The Randles model satisfactorily describes the electrochemical behavior of our system. As expected, systematic increases in resistance describing charge transfer at the electrode are observed with the chemical steps involved in DNA probe attachment. However, anomalous drops in charge-transfer resistance are observed when complementary DNA hybridizes with the probe. We ascribe these to morphological changes in the DNA that facilitate better approach of the redox complexes to weak points in the SAM where charge transfer is efficient. More detailed chemical and spectroscopic studies are needed to understand that phenomenon at a fundamental level. Further improvements in our basic understanding and in optimization of the surface attachment chemistry promise to make this form of electrochemical detection of biomolecular binding at interfaces a powerful analytical tool. Acknowledgment. We are grateful to a grant from Infotonics for this work. We also appreciate the help of Professor Tom Jones and Professor Joe Dinnocenzo in making this work possible. LA048083A

(29) Umek, R. M.; Lin, S. W.; Vielmetter, J.; Terbrueggen, R. H.; Irvine, B.; Yu, C. J.; Kayyem, J. F.; Yowanto, H.; Blackburn, G. F.; Farkas, D. H.; Chen, Y. P. J. Mol. Diagn. 2001, 3, 74-84. (30) Motorola Life Sciences Inc. http://www.motorola.com/ lifesciences/.