Label-free Capacitive Diagnostics: Exploiting Local Redox Probe

Article pubs.acs.org/ac

Label-free Capacitive Diagnostics: Exploiting Local Redox Probe State Occupancy Joshua Lehr,† Flávio C. Bedatty Fernandes,‡ Paulo R. Bueno,*,‡ and Jason J. Davis*,† †

Department of Chemistry, University of Oxford, South Parks Road, Oxford OX1 3QZ, United Kingdom Physical Chemistry Department, Institute of Chemistry, Universidade Estadual Paulista (São Paulo State University), CP 355, 14800-900 Araraquara, São Paulo, Brazil

‡

ABSTRACT: An electrode surface confined redox group contributes to a substantial potential-dependent interfacial charging that can be sensitively probed and frequency-resolved by impedance-derived capacitance spectroscopy. In utilizing the sensitivity of this charging fingerprint to redox group environment, one can seek to generate derived sensory configurations. Exemplified here through the generation of mixed molecular films comprising ferrocene and antibody receptors to two clinically important targets, the label-free methodology is able to report on human prostatic acid phosphatase (PAP), a tumor marker, with a limit of detection of 11 pM and C-reactive protein with a limit of detection of 28 pM. Both assays exhibit linear ranges encompassing those of clinical value. 2).15−17 It was our hypothesis that this charging fingerprint would be very sensitively dependent on the redox probe environment, and thus the integration of redox reporters within mixed films composed also of antibody receptors might underpin biodetection. This approach, which does not require any (potentially perturbative) prelabeling of the receptor (Figure 1), was observed to be effective in preliminary work.18 Cr is spectrally resolved from nonfaradaic film capacitance (reported by Ct and Cm)17 and has an expected highly sensitive potential dependence (see Figure 2).16 The relationship between the faradaic current density (jr) and redox capacitance is given by jf = Crs, where s is the voltage scan rate.16 Therefore, Cr reports directly on redox site occupancy and maps out changes resolvable in classic cyclic voltammetry (CV) analyses.15−17 Cr specifically depends on electrode potential according to f = n/Γ = F(Er, μe), where μe is the electron chemical potential and n is the number of occupied redox centers and is maximal with the electrode poised at the electrochemical half-wave potential. It is possible to demonstrate that this Cr charging fingerprint has inseparable quantum and electrostatic components,19 both of which are sensitive to local environment.18 Since this interfacial charging depends on the quantum mechanical coupling of redox and electrode states, the former spanning a Gaussian in distribution, one can quantify by integrating over all contributing energy levels:

A

n ability to selectively and quantitatively assay biomarkers is critical to biological evaluation and diagnosis, supporting the early detection of infectious disease, cancer, and cardiovascular conditions. Traditional detection methods have relied on enzyme-linked immunosorbent assays (ELISA).1 Though highly sensitive, the rather awkward multistep nature of ELISA, together with the need for labeled antibodies (with potentially high associated background cost and error), has led to prolonged interest in the development of equivalently sensitive, facile, label-free methods such as those based on surface plasmon resonance (SPR),2 quartz crystal microbalance (QCM),3 and electrochemistry.4 Platforms based on the latter can enable highly sensitive and convenient analyte detection with cheap, and highly tailorable, components.5−9 Of the considerable range of amperometric and voltammetric methodologies available, electrochemical impedance spectroscopy offers a great deal when seeking very high levels of sensitivity without target labeling, presynthesis, or the use of sandwich formats necessitating two specific antibodies per target.10−12 The simplest, and most studied, impedance-based diagnostic assays are based on detecting the selective recruitment of targets at the electrode interface by measure of the increased resistance to charge transfer from diffusive redox-active probes to the underlying electrode (i.e., the charge-transfer resistance, Rct). While the technique has been popular in the development of assays capable of target detection in the picomolar concentration range, it does require the pre-addition of high concentrations of redox reporter to the analytical solution and the post utilization of an often poorly understood “equivalent circuit”.10,11,13,14 In recent work we have reported the ability of standard impedance analyses to report on both the dielectric and faradaic activity of electrode-confined molecular films. The latter is frequency-resolved within a redox-capacitive (Cr) term (Figure © 2014 American Chemical Society

Received: November 16, 2013 Accepted: February 3, 2014 Published: February 3, 2014 2559

dx.doi.org/10.1021/ac403727h | Anal. Chem. 2014, 86, 2559−2564

Analytical Chemistry

Article

Figure 1. (a) General surface architectures comprising mixed PEG-anchored antibody and thiolated ferrocene films. (b) Specific surface chemistry for C-reactive protein (CRP) and prostatic acid phosphatase (PAP) receptive interfaces.

occupancy is thus integrated into gr and experimentally reported through Cr. Our hypothesis was that local changes in dielectric constant and/or electrostatics will perturb this charging fingerprint in a detectable manner. If one can encapsulate such an ability within a selective target-recruiting molecular film, one can feasibly generate potent derived biosensors where target detection is transduced without the need for either target (primary or secondary) labeling or solution dosing. Prostate cancer (PCa) is, after lung cancer, the second most destructive form of cancer, affecting some 4 million men in Europe alone.20 The asymptomatic nature of initial growth makes early diagnosis difficult. To date, detection relies heavily on the assaying of a 34 kDa glycoprotein, serum prostatespecific antigen (PSA). These assays alone are not, however, sufficient to be confirmatory of disease (PSA levels are influenced by a multitude of conditions).21 There is, then, considerable need for increased specificity. This may be achieved through the integration of additional marker quantification. Human prostatic acid phosphatase (PAP) is a major phosphatase enzyme of differentiated prostate epithelial cells22,23 biosynthesized by columnar secretory epithelia of the prostate gland.24 Circulating levels of PAP have long been considered as being diagnostically useful; though serum PAP levels are low in healthy individuals, elevation occurs with metastatic PCa and subsequently correlate with disease progression.25−28 In addition, PAP has been demonstrated to be a good prognostic marker for patients with aggressive disease undergoing local therapy who are at high risk for distant relapse.29 In normal healthy individuals, the plasma levels of PAP are on the order of 1−3 ng·mL−1 (10−30 pM), becoming elevated in a manner correlating with the onset of prostate cancer. Traditional methods for PAP detection, based on ELISA, have linear ranges running approximately 3−30 ng· mL−1 (30−300 pM).30 Electroanalytical methods, as noted, commonly have a high innate sensitivity and simplicity that can be effectively married to miniaturized hardware. As such, they constitute arguably the most practical, quantifiable, and scalable of all low-cost diagnostic assessments of protein presence. There are to the best of our knowledge no previous reports of

Figure 2. (a) Generalised depicted of surface architextures comprising mixed PEG and thiolated ferrocene films with an associated Cr capacitative fingerprint. (b) Depiction of the antibody modified interface in which Cr is now sensitised to the recruitment of a specific target.

Cr(μe ) = e 2

∞

df

∫−∞ gr(μe) dμ dμe e

e2 = kBT

∞

∫−∞ gr(μe)f (1 − f ) dμe

(1)

where gr(μe), the redox density of states, can be written as gr (μe ) =

⎡ (E − μ )2 ⎤ 1 r e ⎥ exp⎢ − ⎢⎣ σg 2π 2σg 2 ⎥⎦

(2)

that is, a Gaussian function with a distribution of Er and a σg standard deviation (σg2 variance).15,16 Any change in redox site 2560

dx.doi.org/10.1021/ac403727h | Anal. Chem. 2014, 86, 2559−2564

Analytical Chemistry

Article

Figure 3. Nyquist plots showing the change in capacitive fingerprint as the electrodes are modified; bare gold electrode response (black), PEG− CO2H + ferrocenethiol layer (red), and the latter after (a) anti-CRP and (b) anti-PAP immobilization (blue). Cr is approximated by the diameter of the semicircle (since the bare gold response is much lower). All measurements were performed at the half-wave potential of ferrocene (0.45 V) vs Ag|AgCl, where redox-capacitive response is expected to be maximal.

sequently the electrode was washed with phosphate-buffered saline (PBS, pH 7.4), dried under nitrogen gas, and then immersed in a 5 μM solution of anti-CRP or anti-PAP for 3 h. Finally the anti-PAP electrodes were immersed in 10 μM bovine serum albumin (BSA) in PBS for 30 min to block remaining carboxyl groups, while the anti-CRP interfaces were immersed in 1 M ethanolamine (pH 8.5) in H2O for 20 min. Films were subsequently characterized by cyclic voltammetry, reductive stripping voltammetry, and impedance/capacitance spectroscopies. Electrochemical Impedance Measurements. All electrochemical measurements were undertaken on a Autolab PGSTAT-12 fitted with an FRA2 module. A three-electrode cell setup was used with gold (BASi) working electrode, platinum wire auxiliary electrode, and Ag|AgCl or mercury sulfate reference electrode. Impedance spectra were collected between 1 MHz and 0.01 Hz with amplitude of 10 mV (peak to peak) and were subsequently verified for compliance with Kramers−Kronig linear systems theory by employing the FRA Autolab software. The redox-capacitive nature of these interfaces can be sensitively analyzed by measuring complex Z*(ω) (impedance) function and converting into C*(ω) (capacitance) by C*(ω) = 1/iωZ*(ω), where ω is the angular frequency and i = √−1 (i.e., complex number). Practically, this involves taking the data resolved in a standard impedance analysis [Z*(ω)], sampled across a range of frequencies at any steady-state potential, and converting it phasorially into complex capacitance [C*(ω)] with its real and imaginary components. In processing Z*(ω) data sets in this way, one obtains the imaginary part of the capacitance by noting that C″ = φZ′ and C′ = φZ″, where φ = (ω|Z|2)−1 and |Z| is the modulus of Z*. If one carries out this analysis outside of the surface potential window where redox activity is observed and then inside the potential window, the difference (the “redoxonly capacitive term”, Cr) is obtained.15 Since, in the absence of a redox film, charging capacitance is comparatively very small, Cr can, alternatively, be estimated by simple subtraction of the former (without prior acquisition of capacitance data in and out of the redox potential window for all interfaces). It is notable that this faradaic capacitive component is significantly more sensitive to the interfacial environment, such as binding events, than is nonfaradaic charging (see redox out response in Figure 5, for instance).15 Sensor Testing. Initially, impedance spectra were recorded for the antibody/ferrocene electrodes prior to any incubation in analyte. Interfaces were then placed in a 0.15 M phosphate buffer solution (pH 7.3) containing specific quantifications of CRP or PAP and impedance spectra were recorded after incubation for 30 min, followed by washing in PBS solution

electrochemical assays that have been developed for PAP detection. C-reactive protein (CRP) is an important biomarker for cardiac events and inflammation generally,31 where its quantification can be an indispensable early warning assay. Present clinical CRP detection methods are based on turbidimetric and nephelometric methods,32,33 with human CRP ELISA kits also employed. These approaches, though robust, suffer from limitations associated with sensitivity, cost, selectivity, and processing time,34 and accordingly, a number of alternative label-free methods based on QCM,35 SPR,36,37 piezoelectric cantilevers,38 or electrochemistry have been presented.4,12,31,39,40 Unfortunately, most reports have not been shown to be selective39 or to span the clinically useful range.40,41 Building upon prior impedimetric assays by us,12,31 and others40,42−44 we show herein that a highly selective assay spanning 10 pM−10 nM can be established without the need to add amplifying redox probe to solution. We specifically demonstrate that the capacitive fingerprint of mixed alkylferrocene−PEG−antibody films [where PEG is poly(ethylene glycol)] selectively and sensitively transduces the interfacial binding of antibody-specific targets.

■

EXPERIMENTAL SECTION

Electrode Preparation. Gold electrode disks (2.0 mm diameter, from Metrohm, or 1.6 mm, from BASi) were mechanically polished with aluminum oxide pads or diamond spray on polishing cloth (Kemet) of progressively decreasing particle size: 1, 0.3, and 0.05 μm, with intermittent sonication in water. The electrodes were then electrochemically polished in a deaerated NaOH or KOH solution (0.5 mol·L−1) between the potentials −1.5 and −0.5 V versus Ag|AgCl or Ag wire at a scan rate of 100 mV·s−1 and then in deaerated 0.5 M H2SO4 between −0.2 and 1.5 V at 100 mV·s−1 until the gold reduction peak was stabilized (around 50 cycles). Electroactive areas were evaluated by integration of the cathodic peak from gold electropolishing voltammograms and converted to the real surface area by use of a conversion factor of 400 μC·cm−2.45 These determinations (0.033−0.04 cm2) were used in the normalization of absolute recorded capacitance. Film Preparation. Self-assembled monolayers (SAMs) were prepared by immersion of the gold electrodes (BASi) in a mixed (2 mM) solution of thiol−PEG−CO2H (Prochimia) and 11-ferrocenylundecanethiol (Sigma Aldrich) (1:100). The former provides a protein-resistant surface46 that is, additionally, antibody biocompatible.47,48 The PEG carboxy terminations were activated with an aqueous solution containing 0.1 M N-hydroxysuccinimide (NHS) and 0.4 M 1-ethyl-3-(3dimethylaminopropyl)carbodiimide (EDC) for 40 min. Sub2561

dx.doi.org/10.1021/ac403727h | Anal. Chem. 2014, 86, 2559−2564

Analytical Chemistry

Article

Figure 4. (a, c) Impedimetric and capacitive (b, d) Nyquist plots for redox-active interfaces, either (a, b) CRP-receptive or (c, d) PAP-receptive, after incubation in various concentrations of target. The higher responsiveness exhibited by the capacitive plots, in comparison to the impedimetric plots, underlines the utility of capacitive redox biosensing.

before electrochemical analysis. After calculation of the relevant Cr values, the data was fitted to a calibration curve by use of 1/ ΔCr, calculated from 1/Cr − 1/Cr,initial, where 1/Cr,initial is the inverse of the capacitance of the antibody interface prior to any addition of target. The limit of detection (LOD)49 was determined as previously reported according to IUPAC standardization. KD values were obtained by fitting a Langmuir isotherm to obtained analytical curves (see Figure 6) by use of the relationship [target]/Cr as a function of [target].12 In evaluating interfacial specificity, streptavidin and the blood proteins fetuin A and BSA were used as controls. Fetuin A is a glycoprotein formed by liver cells and secreted into the serum in high concentration.50,51

■

RESULTS AND DISCUSSION Characterization of Interface Construction. Construction of the target recruiting electroactive interfaces was monitored by capacitance spectroscopy (Figure 3). Capacitive Nyquist plots were specifically acquired (at the half-wave potential of the ferrocene/ferrocenium couple, Fc/Fc+) for bare gold, ferrocene/PEG SAMs, and the antibody-modified interfaces. An increase in capacitance from double-layer values of ∼10 to ∼300 μF·cm−2 is observed upon modification of the bare electrode with Fc/PEG films, attributed to the introduction of redox capacitance associated with the ferrocene groups. Subsequent coupling of the antibody to the activated surface resulted in a notable decrease in capacitance to ∼250 μF·cm−2, attributed to a perturbation of the density of states (DOS) occupation (see eqs 1 and 2) associated with environmental change around the ferrocene redox centers.19 Redox-Capacitance Response of Target Binding. Once the faradaic charging fingerprints of these films had been established, their selective response to target was analyzed. As noted, the capacitance spectroscopy method utilized here is distinct from traditional impedance-based assays where resistance to charge transfer would be expected and tracked through changes in Nyquist diagrams of complex impedance (Figure 4a,c). As shown below, the impedance fingerprints of these interfaces are neither sensitive to target capture nor such that charge-transfer resistance can be reliably estimated by fitting to the equivalent circuit of Figure 2b. In addition to the capacitive Nyquist plots shown in Figure 4, acquired data can be represented in the form of capacitive Bode plots, in which the frequency-resolved influence of ferrocene charging on real and imaginary capacitive elements is resolved (Figure 5). Bode plots of real capacitance depict redox charging

Figure 5. (a, c) Real and imaginary (b, d) capacitive Bode plots for redox-active interfaces, either (a, b) CRP-receptive or (c, d) PAPreceptive, after incubation with specific target concentrations. The real component shows the redox capacitance as a charging plateau, observed as a peak at a frequency corresponding to the rate of electron transfer in the analogous imaginary component. Only very low capacitances are observed outside the redox window of the ferrocene/ ferrocenium couple, indicating, as noted in the main text, that nonfaradaic charging is much smaller than faradaic charging in these films. The “redox in” signals were recorded at the half-wave potential of the ferrocene (∼0.45 vs Ag|AgCl) with “redox out” (nonfaradaic) signals acquired at −0.1 V vs Ag|AgCl. The frequency axis (logarithmic scale) is reported here as the base 10 exponents only for clarity.

as a charging plateau, while the imaginary component is observed as a peak with a maximum at a frequency that correlates directly to the rate of electron transfer between the surface-bound ferrocene and the gold substrate, as reported previously.15 Importantly, when data are recorded outside the “redox window” of the ferrocene, the capacitive charging is both very small and largely unresponsive to target recognition, with typically less than 2% of the response that is observed within the faradaic window (see Figure 5). Faradaic capacitive charging related to the DOS occupancy is turned on by 2562

dx.doi.org/10.1021/ac403727h | Anal. Chem. 2014, 86, 2559−2564

Analytical Chemistry

Article

Figure 6. Analytical capacitive curves reporting Cr for (a) CRP detection at CRP receptive interfaces and PAP detection at PAP-receptive interfaces (b). Plotted data points represent means and standard deviations across three measured interfaces in each case. Errors are reported at 95% confidence. R2, the coefficient of determination, for CRP is 0.994 and R2 for PAP is 0.990.

Figure 7. Redox-capacitive response of (a) anti-CRP and (b) anti-PAP interfaces in the presence of nonspecific protein as a percentage of the response (Rn) seen upon selective binding. Percent response (ordinate axis), Rn = 100(Rn − R0)/R0, where calculated with respect to CRP- and PAPspecific interface response for zeroed target concentration (R0). Across both interfaces, nonspecific responses were typically less than 4% of R0, in contrast to specific responses of 48−50% of R0.

lization of anti-CRP/PAP and ferrocene capacitative reporters on gold substrates. Both interfaces exhibited excellent selectivity and sensitivity, with substantial and clinically useful linear ranges. For both targets, detection limits exceed those of any prior label-free electroanalysis. The perturbations of interfacial redox capacitance in these label-free assays will depend on antibody binding efficacy, antibody surface density, binding constant, and the size and dielectric character of the target (one might expect, as indeed is observed here, that the binding of a 115 kDa pentamer, CRP, generates a greater interfacial perturbation that the binding of monomeric 50 kDa PAP). The results presented here underline the utility and considerable potential of the use of capacitance spectroscopy in the development of highly sensitive and selective label-free sensors capable of detecting a wide range of biological makers. The methodology is simple, frequency-optimized (and therefore readily applicable within multiplexing formats), does not rely on any modeling/fitting or associated assumptions for data analysis, and required no prior doping of analytical solution with redox probe or label.

moving the electrode poise into the redox window and responds in a highly sensitive manner to target. KD determinations were 390 ± 170 and 278.0 ± 2.65 pM for the anti-CRP and anti-PAP interfaces respectively. Both are somewhat lower than analogous SPR determinations with single-component films,37 an observation potentially reflective of the advantages associated with high levels of antibody surface dilution used here. Target Quantification. It is evident from Figures 4 and 5 that Cr decreases in response to target presence for both antiCRP and anti-PAP interfaces. These responses are specific to these targets (Figure 7). Linear responses are obtained for both systems by plotting the inverse of Cr (normalized for surface area) against the natural logarithm of concentration (Figure 6). The linear range of the CRP sensor spans from 50 pM to 100 nM, aligning well with the values that are clinically useful.44,52 The limit of detection (LOD) was determined to be 28 pM,53 the lowest of any reported clinically useful impedimetric CRP sensor.12,31,41,44 For the PAP-receptive interfaces, a LOD of 11 pM and a linear range of 50 pM−10 nM are determined, comparing favorably with ELISA analyses30 and the clinically useful range, given that healthy plasma levels of PAP are on the order of 10− 30 pM and elevated levels around 100 pM are indicative of prostatic cancer.30

■

AUTHOR INFORMATION

Corresponding Authors

*Phone +55 16 3301 9642; fax +55 16 3322 2308; e-mail [email protected]. *Phone +44 18 6527 5914; e-mail [email protected].

■

CONCLUSIONS We have herein reported the fabrication of PAP and CRP redox-capacitance biosensing interfaces by covalent coimmobi-

Notes

The authors declare no competing financial interest. 2563

dx.doi.org/10.1021/ac403727h | Anal. Chem. 2014, 86, 2559−2564

Analytical Chemistry

■

Article

(33) Roberts, W. L.; Moulton, L.; Law, T. C.; Farrow, G.; CooperAnderson, M.; Savory, J.; Rifai, N. Clin. Chem. 2001, 47, 418−425. (34) Pearson, T. A.; Mensah, G. A.; Hong, Y. L.; Smith, S. C. Circulation 2004, 110, E543−E544. (35) Kim, N.; Kim, D.-K.; Cho, Y.-J. Sens. Actuators, B 2009, 143, 444−448. (36) Hu, W. P.; Hsu, H. Y.; Chiou, A.; Tseng, K. Y.; Lin, H. Y.; Chang, G. L.; Chen, S. J. Biosens. Bioelectron. 2006, 21, 1631−1637. (37) Meyer, M. H. F.; Hartmann, M.; Keusgen, M. Biosens. Bioelectron. 2006, 21, 1987−1990. (38) Wee, K. W.; Kang, G. Y.; Park, J.; Kang, J. Y.; Yoon, D. S.; Park, J. H.; Kim, T. S. Biosens. Bioelectron. 2005, 20, 1932−1938. (39) Hennessey, H.; Afara, N.; Omanovic, S.; Padjen, A. L. Anal. Chim. Acta 2009, 643, 45−53. (40) Qureshi, A.; Gurbuz, Y.; Kallempudi, S.; Niazi, J. H. Phys. Chem. Chem. Phys. 2010, 12, 9176−9182. (41) Chen, X.; Wang, Y.; Zhou, J.; Yan, W.; Li, X.; Zhu, J.-J. Anal. Chem. 2008, 80, 2133−2140. (42) Buch, M.; Rishpon, J. Electroanalysis 2008, 20, 2592−2594. (43) Centi, S.; Sanmartin, L. B.; Tombelli, S.; Palchetti, I.; Mascini, M. Electroanalysis 2009, 21, 1309−1315. (44) Vermeeren, V.; Grieten, L.; Vanden Bon, N.; Bijnens, N.; Wenmackers, S.; Janssens, S. D.; Haenen, K.; Wagner, P.; Michiels, L. Sens. Actuators, B 2011, 157, 130−138. (45) Trasatti, S.; Petrii, O. A. Pure Appl. Chem. 1991, 63, 711−734. (46) Alcantar, N. A.; Aydil, E. S.; Israelachvili, J. N. J. Biomed. Mater. Res. 2000, 51, 343−351. (47) Cho, H. Y.; Srinivasan, A.; Hong, J.; Hsu, E.; Liu, S.; Shrivats, A.; Kwak, D.; Bohaty, A. K.; Paik, H.-j.; Hollinger, J. O.; Matyjaszewski, K. Biomacromolecules 2011, 12, 3478−3486. (48) Klapshina, L. G.; Douglas, W. E.; Grigoryev, I. S.; Ladilina, E. Y.; Shirmanova, M. V.; Mysyagin, S. A.; Balalaeva, I. V.; Zagaynova, E. V. Chem. Commun. 2010, 46, 8398−8400. (49) Long, G. L.; Winefordner, J. D. Anal. Chem. 1983, 55, A712. (50) Pedersen, K. O. Nature 1944, 154, 575−575. (51) Singh, M.; Sharma, P. K.; Garg, V. K.; Mondal, S. C.; Singh, A. K.; Kumar, N. J. Pharm. Pharmacol. 2012, 64, 1703−1708. (52) Pepys, M. B. In Encyclopedia of Immunology, 2nd ed.; Peter, J. D., Ed.; Elsevier: Oxford, U.K., 1998. (53) The LOD of the CRP sensor reported herein is significantly lower than that reported in a preliminary study.18 The previous surface constructs consisted of antibody physisorbed on a mixed ferrocenylundecanethiol/alkanethiol film. Herein, antibodies have been covalently immobilized on a mixed ferrocenylundecanethiol/PEG interface. The approximate order of magnitude improvement in assay sensitivity is in line with expectations for an interface that is chemically different and both more biocompatible and appreciably more nonfouling.

ACKNOWLEDGMENTS We acknowledge São Paulo state research funding agency (FAPESP) and CNPq. F.C.B.F. acknowledges CAPES for his Ph.D. scholarship. We additionally thank Márcio Santos for preliminary discussions about the PAP assays.

■

REFERENCES

(1) Zangar, R. C.; Daly, D. S.; White, A. M. Expert Rev. Proteomics 2006, 3, 37−44. (2) Reddy, P. J.; Sadhu, S.; Ray, S.; Srivastava, S. Clin. Lab. Med. 2012, 32, 47−72. (3) Pavey, K. D. Expert Rev. Mol. Diagn. 2002, 2, 173−186. (4) Luo, X.; Davis, J. J. Chem. Soc. Rev. 2013, 42, 5944−5962. (5) Berggren, C.; Bjarnason, B.; Johansson, G. Electroanalysis 2001, 13, 173−180. (6) Davis, J. J.; Tkac, J. In Engineering the Bioelectronic Interface; Davis, J. J., Ed.; Royal Society of Chemistry: London, 2009. (7) Estrela, P.; Paul, D.; Song, Q.; Stadler, L. K. J.; Wang, L.; Huq, E.; Davis, J. J.; Ferrigno, P. K.; Migliorato, P. Anal. Chem. 2010, 82, 3531− 3536. (8) White, R. J.; Kallewaard, H. M.; Hsieh, W.; Patterson, A. S.; Kasehagen, J. B.; Cash, K. J.; Uzawa, T.; Soh, H. T.; Plaxco, K. W. Anal. Chem. 2012, 84, 1098−1103. (9) Xu, M. Y.; Lou, X. L.; Davis, J. J. Biosens. Bioelectron. 2013, 39, 21−25. (10) Lisdat, F.; Schaefer, D. Anal. Bioanal. Chem. 2008, 391, 1555− 1567. (11) Daniels, J. S.; Pourmand, N. Electroanalysis 2007, 19, 1239− 1257. (12) Johnson, A.; Song, Q.; Ferrigno, P. K.; Bueno, P. R.; Davis, J. J. Anal. Chem. 2012, 84, 6553−6560. (13) Rodriguez, M. C.; Kawde, A. N.; Wang, J. Chem. Commun. 2005, 4267−4269. (14) Bogomolova, A.; Komarova, E.; Reber, K.; Gerasimov, T.; Yavuz, O.; Bhatt, S.; Aldissi, M. Anal. Chem. 2009, 81, 3944−3949. (15) Bueno, P. R.; Mizzon, G.; Davis, J. J. J. Phys. Chem. B 2012, 116, 8822−8829. (16) Bueno, P. R.; Fabregat-Santiago, F.; Davis, J. J. Anal. Chem. 2013, 85, 411−417. (17) Goes, M. S.; Rahman, H.; Ryall, J.; Davis, J. J.; Bueno, P. R. Langmuir 2012, 28, 9689−9699. (18) Fernandes, F. C. B.; Góes, M. S.; Davis, J. J.; Bueno, P. R. Biosens. Bioelectron. 2013, 50, 437−440. (19) Bueno, P. R.; Davis, J. J. Anal. Chem. 2014, 86, 1337−1341. (20) Figg, W. D., II; Figg, W. D., Sr. Cancer Biol. Ther. 2010, 10, 1233−1234. (21) Qu, Y. S.; Adam, B. L.; Yasui, T.; Ward, M. D.; Cazares, L. H.; Schellhammer, P. F.; Feng, Z. D.; Semmes, O. J.; Wright, G. L. Clin. Chem. 2002, 48, 1835−1843. (22) Bodansky, O. Cancer 1974, 33, 364−370. (23) Azumi, N.; Traweek, S. T.; Battifora, H. Am. J. Surg. Pathol. 1991, 15, 785−790. (24) Vihko, P. Clin. Chem. 1978, 24, 1783−1787. (25) Ahmann, F. R.; Schifman, R. B. J. Urol. 1987, 137, 431−434. (26) Robles, J. M.; Morell, A. R.; Redorta, J. P.; Mateos, J. A. D.; Rosello, A. S. Eur. Urol. 1988, 14, 360−366. (27) Simsek, U.; Kutlu, S.; Yavascaouglu, I.; Oktay, B.; Ozyurt, M. Eur. Urol. 1992, 21, 111−114. (28) Muniyan, S.; Chaturvedi, N. K.; Dwyer, J. G.; LaGrange, C. A.; Chaney, W. G.; Lin, M.-F. Int. J. Mol. Sci. 2013, 14, 10438−10464. (29) Taira, A. L.; Merrick, G.; Wallner, K.; Dattoli, M. Oncology (New York) 2007, 21, 1003−1010. (30) Vandieijenvisser, M. P.; Delaere, K. P. J.; Gijzen, A. H. J.; Brombacher, P. J. Clin. Chim. Acta 1988, 174, 131−140. (31) Bryan, T.; Luo, X.; Bueno, P. R.; Davis, J. J. Biosens. Bioelectron. 2013, 39, 94−98. (32) Roberts, W. L.; Sedrick, R.; Moulton, L.; Spencer, A.; Rifai, N. Clin. Chem. 2000, 46, 461−468. 2564

dx.doi.org/10.1021/ac403727h | Anal. Chem. 2014, 86, 2559−2564