Labeless Immunosensor Assay for Prostate Specific Antigen with

Jul 22, 2008 - Microelectrode arrays were interrogated using ac impedance protocols before and following exposure to PSA solutions. Our preliminary re...
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Anal. Chem. 2008, 80, 6198–6205

Labeless Immunosensor Assay for Prostate Specific Antigen with Picogram per Milliliter Limits of Detection Based upon an ac Impedance Protocol Andrew C. Barton, Frank Davis, and Se ´ amus P. J. Higson* Cranfield Health, Cranfield University, Silsoe, Beds, MK45 4DT, U.K. This paper describes the development of labeless immunosensors for the prostate cancer marker prostate specific antigen (PSA). Poly(1,2-diaminobenzene) was electrodeposited onto screen-printed carbon electrodes, and this modified surface was sonochemically ablated to form a microelectrode array. Polyaniline was electropolymerized within these pores to form a microarray of conductive polyaniline protrusions. Two methods were utilized to immobilize antibodies for prostate specific antigen (APSA). The first involved entrapment of APSA during electropolymerization of the polyaniline. The second utilized a polyaniline array as a substrate to immobilize a biotinylated APSA using a classical avidin-biotin affinity approach. Microelectrode arrays were interrogated using ac impedance protocols before and following exposure to PSA solutions. Our preliminary results show that concentration/ac response relationships were recorded over very different ranges; sensors fabricated using an affinity approach exhibited detection limits 1000 times lower than those formulated by the entrapment method. This demonstrates that assembly protocols have major effects on immunosensor performance. The principle of immunoassays established in 1959 led to the development of the widely used radioimmunoassay.1 Later, unconnected work2 in 1962 pioneered the concept of a biosensor, exploiting the selectivity of enzymes for analytical purposes via immobilizing enzymes on the surface of electrochemical sensors and measuring the results of enzymatically catalyzed reactions. Incorporation of antibodies into conducting polymer films was first reported3 in 1991. Pyrrole was polymerized onto a platinum wire substrate from solutions containing antihuman serum albumin (anti-HSA), allowing antibody incorporation into the polypyrrole film. The resultant electrode was found to respond specifically to the antigen. Other early work developed immunosensors for other analytes, including p-cresol,4 thaumatin,5 and polychlorinated * Corresponding author. Fax (+44) 1234 758370, e-mail s.p.j.higson@ cranfield.ac.uk. (1) Yalow, R. S.; Berson, S. A. Nature 1959, 184, 1648–1649. (2) Clark, L. C.; Lyons, I. R. Ann. N.Y. Acad. Sci. 1962, 102, 29. (3) John, R.; Spencer, M.; Wallace, G. G.; Smyth, M. R. Anal. Chim. Acta 1999, 249, 381–385. (4) Barnett, D.; Laing, D. G.; Skopec, S.; Sadik, O. A.; Wallace, G. G. Anal. Lett. 1994, 27, 2417.

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biphenyls.6 There has over the past few years been a burgeoning interest in the development of electrochemical immunosensors, as detailed in several recent reviews.7–9 We have previously shown that up to 2-3 µg of antibodies for BSA and digoxin may be successfully entrapped into conducting polypyrrole films.10 Further work utilizing an ac impedance protocol11 led to the development of immunosensors for digoxin and bovine serum albumin. Our later work utilized polyaniline coated screen-printed carbon electrodes as substrates for antibody immobilization via avidin-biotin interactions. This enabled the construction of immunosensors for ciprofloxacin12 and myelin basic protein, a marker for stroke and multiple sclerosis.13 Our group has pioneered the development of sonochemically fabricated microarrays of conductive polymers,14 the schematic for the formation of which is shown within Figure 1a. Poly(1,2diaminobenzene) can be electrodeposited on a variety of conductive surfaces to form a self-limiting insulating layer of highly reproducible thickness (50-70 nm).15 We have utilized commercial screen printed three electrode strips, electrochemically coating the working electrodes with an insulating polymer of 1,2diaminobenzene.15 Sonochemical ablation is then used to ablate or “drill” holes in this insulating material16 with diameters of 0.1-3 µm and a density of up to 120 000 pores cm-2. The arrays may be used as substrates for further electropolymerization reactions, generating arrays of conducting polyaniline protrusions, potentially (5) Sadik, O. A.; John, M. J.; Wallace, G. G.; Barnett, D.; Clarke, C.; Laing, D. G. Analyst 1994, 119, 1997–2000. (6) Bender, S.; Sadik, O. A. Environ. Sci. Technol. 1998, 32, 788–797. (7) Rodriguez-Mozaz, S.; de Alda, M. J. L.; Barcelo, D. Anal. Bioanal. Chem. 2006, 386, 1025–1041. (8) Diaz-Gonzalez, M.; Gonzalez-Garcia, M. B.; Costa-Garcia, A. Electroanalysis 2005, 17, 1901–1918. (9) Cosnier, S. Electroanalysis 2005, 17, 1701–1715. (10) Grant, S.; Davis, F.; Pritchard, J. A.; Law, K. A.; Higson, S. P. J.; Gibson, T. D. Anal. Chim. Acta 2003, 495, 21–32. (11) Grant, S.; Davis, F.; Law, K. A.; Barton, A. C.; Collyer, S. D.; Higson, S. P. J.; Gibson, T. D. Anal. Chim. Acta 2003, 537, 163–168. (12) Garifallou, G. Z.; Tsekenis, G.; Davis, F.; Millner, P. A.; Pinacho, D. G.; Sanchez-Baeza, F.; Marco, M.-P.; Gibson, T. D.; Higson, S. P. J. Anal. Lett. 2007, 40, 1412–1442. (13) Tsekenis, G.; Garifallou, G. Z.; Davis, F.; Millner, P. A.; Gibson, T. D.; Higson, S. P. J. Anal. Chem. 2008, 80, 2058–2062. (14) Higson, S. P. J. Sensor. International Patent: PCT/GB96/0092, 1996. (15) Myler, S.; Eaton, S.; Higson, S. P. J. Anal. Chim. Acta 1998, 357, 55–61. (16) Davis, F.; Collyer, S. D.; Gornall, D. D.; Law, K. A.; Mills, D. W.; Higson, S. P. J. Chim. Oggi (Chemistry Today) 2007, 25, 28–31. 10.1021/ac800491m CCC: $40.75  2008 American Chemical Society Published on Web 07/22/2008

Figure 1. Formation of polyaniline microarrays: (a) deposition of insulating layer, (b) sonochemical formation of pores, (c) polymerization of aniline, (d) schematic of antibody modified electrodes, and (e) commercial screen-printed electrodes used within this work.

containing entrapped biological species.17 We have utilized microarrays containing entrapped enzymes for the amperometric detection of glucose,17,18 alcohol,19 and a range of organophosphate pesticides20,21 with extremely low limits of detection (10-17 M). Prostate cancer is a disease most frequently encountered in men over fifty and second only to lung cancer for the number of male deaths in the U.S.22 (30 350 deaths in 2005) and U.K.23 (10 000 deaths in 2005). Prostate specific antigen (PSA) is a 34 kDa glycoprotein manufactured almost exclusively by the prostate gland. Normal levels of PSA are below 4 ng mL-1 in serum24 and are often (but not always)25 elevated in the presence of prostate cancer and other prostate disorders. A blood test to measure PSA is the most effective test currently available for the early detection of prostate (17) Barton, A. C.; Collyer, S. D.; Davis, F.; Gornall, D. D.; Law, K. A.; Lawrence, E. C. D.; Mills, D. W.; Myler, S.; Pritchard, J. A.; Thompson, M.; Higson, S. P. J. Biosens. Bioelectron. 2004, 20, 328–337. (18) Myler, S.; Davis, F.; Collyer, S. D.; Higson, S. P. J. Biosens. Bioelectron. 2004, 20, 408–412. (19) Myler, S.; Davis, F.; Collyer, S. D.; Gornall, D. D.; Higson, S. P. J. Biosens. Bioelectron. 2005, 21, 666–671. (20) Pritchard, J. A.; Law, K. A.; Vakurov, A.; Millner, P. A.; Higson, S. P. J. Biosens. Bioelectron. 2004, 20, 765772) (21) Law, K. A.; Higson, S. P. J. Biosens. Bioelectron. 2005, 21, 1914–1924. (22) Jemal, A.; Murray, T.; Ward, E.; Samuels, A.; Tiwari, R. C.; Ghafoor, A.; Feuer, E. J.; Thun, M. J. Ca-Cancer J. Clin. 2005, 55, 10–30. (23) Office for National Statistics Mortality Statistics: Cause, 2005. Series DH2, no. 32, 2005. (24) Wu, J.; Fu, Z.; Yan, F.; Ju, H. Trends Anal. Chem. 2007, 26, 679–688.

cancer. Higher than normal levels of PSA are associated with both localized and metastatic prostate cancer. Patient recovery is enhanced by early detection of prostate cancer, especially at the organ-confined stage. This highlights the need for reliable diagnostic tests for the rapid determination of both free PSA and total PSA. A range of immunosensors for PSA have already been developed as detailed in a recent review.24 For example, antibodies to PSA were immobilized at the surface of a pH sensitive polymer.26,27 Binding of antigen led to localized increases in pH, causing polymer degradation, which was monitored by ac impedance, allowing detection of PSA with a limit of 3 ng mL-1. Other workers have immobilized anti-PSA onto piezoresistive microcantilevers28 and incorporated these within microfluidic flow cells to give devices capable of detection of 10 ng mL-1 PSA. Microcantilevers of different geometries were found to be capable of detecting PSA between 0.2 ng mL-1 and 60 µg mL-1 in a background of HSA and human plasminogen.29 Surface (25) Thompson, I.; Pauler, D. K.; Goodman, P. J.; Tangen, C. M.; Lucia, M. S.; Parnes, H. L.; Minasian, L. M.; Ford, L. G.; Lippman, S. M.; Crawford, E. D.; Crowley, J. J.; Coltman, C. A. N. Engl. J. Med. 2004, 350, 2239–2246. (26) Fernandez-Sanchez, C.; McNeil, C. J.; Rawson, K.; Nilsson, O. Anal. Chem. 2004, 76, 5649–5656. (27) Fernandez-Sanchez, C.; Gallardo-Soto, A. M.; Rawson, K.; Nilsson, O.; McNeil, C. J. Electrochem. Commun. 2004, 6, 138–143. (28) Wee, K. W.; Kang, G. H.; Park, J.; Kang, J. Y.; Yoon, D. S.; Park, J. H.; Kim, T. S. Biosens. Bioelectron. 2005, 20, 1932–1938. (29) Wu, G. H.; Datar, R. H.; Hansen, K. M.; Thundat, T.; Cote, R. J.; Majumdar, A. Nat. Biotechnol. 2001, 19, 856–860.

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plasmon resonance studies30 of anti-PSA attached to gold surfaces allowed direct detection of 10 ng mL-1 PSA. Colloidal gold31 could be used to amplify the response, allowing detection limits to be lowered to 0.15 ng mL-1. One disadvantage of the entrapment method for antibody immobilization is that microelectrodes are still relatively large compared with any entrapped biomolecules. Should the antigen be too large to enter the polyaniline matrix, obviously only antibodies located at the surface of the protrusion and suitably orientated will be available for antigen binding. Other work comparing monolayers of randomly and specifically orientated antibody fragments32 showed that immunosensor responses double when the fragment is specifically orientated. We therefore have within this work developed labeless immunosensors for PSA using two different methods, with the antibody either entrapped within the polyaniline matrix or alternatively specifically immobilized at the surface of the protrusion, in an attempt to determine whether combining controlled immobilization with the use of our microarray technology would enhance immunosensor sensitivity. The sensors were based on commercial screen-printed carbon electrodes modified by deposition of a insulating poly(1,2diaminobenzene) film followed by sonochemical ablation to produce microelectrode arrays. Conductive polyaniline protrusions were grown from these pores, either containing entrapped antibodies or alternatively chemically modified with a commercial biotinylating reagent. Neutravidin was complexed with the immobilized biotin to generate a surface suitable for the further binding of biotinylated selective antibodies to PSA as described earlier12,13 and as shown schematically in Figure 1d. A nonspecific IgG antibody was incorporated instead of the anti-PSA to fabricate control electrodes. This permitted subtraction of unspecific interactions from the specific binding response of PSA. This approach helps to increase the stability and reliability of these sensors when applied to clinical samples. Electrodes constructed via the entrapment method are capable of detecting the antigen within the required physiological range whereas those constructed using affinity interactions have limits of detection (3 times the standard deviation of the baseline value) up to 3 orders of magnitude lower. EXPERIMENTAL SECTION Materials and Equipment. Sourcing of chemicals and equipment and preparation of buffer solutions is described in detail within the Supporting Information. Commercial screen-printed carbon electrodes (Figure 1e) containing carbon working and counter electrodes and an Ag/AgCl reference electrode were obtained from Microarray Ltd., Manchester, U.K. The surface area of the working electrode was 0.2178 cm2. For antibody biotinylation, the procedure outlined in the BK101 kit was followed (see manufacturers instructions for details). Antibodies against PSA were reconstituted in pH 7.4 phosphate buffer. Biotinylated antibodies were kept frozen in aliquots of 200 µL at a concentration of 1 mg mL-1 until required. (30) Huang, L.; Reekmans, G.; Saarens, D.; Freid, J. M.; Frederix, F.; Francis, L.; Muyldermans, S.; Campitelli, A.; Van Hoof, C. Biosens. Bioelectron. 2005, 21, 483–490. (31) Besselink, G. A. J.; Kooyman, R. P. H.; van Os, P. J. H. J.; Engbers, G. H. M.; Schasfoort, R. B. M. Anal. Biochem. 2004, 333, 165–173. (32) Bonroy, K.; Federix, F.; Reekmans, G.; Dewolf, E.; Palma, R. D.; Borgha, G.; Declerck, P.; Goddeeris, B. J. Immunol. Methods 2006, 312, 167–181.

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Formation of Microarrays. Sonochemically fabricated microarrays were constructed as previously described,17 and detailed within the Supporting Information. Insulating layers of poly(1,2diaminobenzene) were deposited, ablated with ultrasound, and used as the substrate for the growth of microarrays of conductive polyaniline protusions. Construction of Antibody Microarrays via Entrapment and Affinity Protocols. After deposition of a microarray of conductive polyaniline protrusions, a typical avidin-biotin approach was used to immobilize a biotinylated antibody on the electrode as described previously12,13 and detailed within the Supporting Information. Alternatively a 0.2 M aniline hydrochloride solution was prepared in a pH 4.0 acetate buffer and monoclonal antibody receptor incorporated into the buffered monomer solution prior to polymerization at a resultant concentration of 0.5 mg mL-1. Electrochemical deposition of the polyaniline was performed as detailed within the Supporting Information, leading to formation of polyaniline protrusions containing entrapped antibodies as previously described.17–21 Determination of Antigen Concentration. Alternating current impedance measurements were performed using an ACM Auto ac DSP frequency response analyzer. Following immobilization of antibody, impedance analyses were performed from 1 Hz to 10 000 Hz (±5 mV amplitude perturbation) in pH 7.4 phosphate buffer, i.e., containing no antigen, as a baseline trace. This buffer solution did, however, contain a 50:50 mixture of [Fe(CN)6]3-/4at a concentration of 10 mmol L-1 as redox mediator so as to perform faradaic impedance spectroscopy. The potential of the electrochemical cell is offset to the formal potential of the redox probe (+0.12 V vs Ag/AgCl identified via cyclic voltammetry). Subsequently, APSA doped sensors are exposed to a range of concentrations from 1 to 300 ng mL-1 or pg mL-1 PSA in pH 7.4 phosphate buffer containing [Fe(CN)6]3-/4-. RESULTS AND DISCUSSION Impedance Profiles of the Electrodes. Figure 2a shows the Nyquist curves obtained for specific response in the concentration range of 0-10 ng mL-1 PSA for a PANI modified, carbon microelectrode array sensor assembly with APSA immobilized via an entrapment method. The Z′ (real) component of the impedance increases steadily with decreasing frequency whereas the Z′′ (imaginary) component increases to a maximum value before falling as the frequency approaches 1 Hz. This type of impedance spectrum is indicative of a surface-modified electrode system where the electron transfer is slow and the impedance is controlled by the interfacial electron transfer.33 Similar experiments were carried out utilizing an APSA microarray formulated by the affinity approach. Initial exposures to PSA as described above showed saturation at 1 ng mL-1; no increases in response were observed for higher concentrations. Therefore a range of more dilute solutions (1-300 pg mL-1) were utilized. Figure 2b shows a Nyquist plot obtained for these electrodes exposed to 0-10 pg mL-1 PSA. These results show the affinity based electrodes to be capable of detecting much lower levels of the antigen. Computer fitting of experimental data to a theoretical model, represented by a simple equivalent circuit, is performed by (33) Katz, E.; Willner, I. Electroanalysis 2003, 15, 913–947.

Figure 2. Specific binding of PSA for (a) entrapped APSA, Nyquist plot, concentrations of PSA (A) 0; (B) 1; (C) 2; (D) 3; (E) 4; (F) 5; and (G) 10 ng mL-1 and (b) affinity immobilized APSA, Nyquist plot, concentrations of PSA (A) 0; (B) 1; (C) 2; (D) 3; (E) 4; (F) 5, and (G) 10 pg mL-1.

dedicated ACM software. For a more thorough evaluation of the data obtained, changes in the Nyquist curves may be translated into electron transfer resistance changes. We compared experimentally determined and simulated values of electron transfer impedance and found a 100% correlation to two decimal places. The relative electron transfer resistance changes from the baseline response at each concentration are plotted in Figure 3a,b. Entrapped APSA immunosensors gave linear responses to the analyte from 1 ng mL-1 to 200 ng mL-1 PSA. Affinity based immunosensors gave a linear response from 1 pg mL-1 to 100 pg mL-1 PSA. The insulation of the modified electrode upon formation of stable antibody-antigen immunocomplexes hinders the electron transfer kinetics of the redox probe resulting in the increase of electron transfer resistance. The electron transfer resistance increases with increasing antigen concentration for both types of electrode. Limits of detection (3 times the standard deviation of the baseline value) were 1 ng mL-1 (entrapped APSA) or 1 pg

mL-1 (affinity-bound APSA). The changes tended toward a plateau above 200 ng mL-1 (entrapped) or 100 pg mL-1 (affinity), suggesting that the immunosensors approach saturation. Nonspecific interactions can interfere with immunosensor performance, leading to erroneously elevated results. Identical sets of immunosensors were therefore fabricated by both methods utilizing a nonspecific IgG antibody (AIgG) in place of the specific APSA antibody. Results for these electrodes were obtained in exactly the same manner as for the specific electrodes. Nyquist plots for these systems (not shown) demonstrate that while binding of PSA to the nonspecific immunosensors occurs, responses are much smaller than for APSA modified sensors. Calibration plots for all these measurements can also be drawn (Figure 3c,d) and demonstrate much lower responses for the nonspecific antibody. A linear nonspecific response is observed for affinity based AIgG immunosensors from 1 to 100 pg mL-1 PSA with a plateau above this concentration; similar behavior occurs for entrapped AIgG immunosensors between 1 and 100 Analytical Chemistry, Vol. 80, No. 16, August 15, 2008

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Figure 3. Calibration plot showing changes in electron transfer resistance vs PSA concentration: (a) entrapped APSA (0-300 ng mL-1), (b) affinity-immobilized APSA (0-300 pg mL-1), (c) entrapped AIgG (0-300 ng mL-1), and (d) affinity-immobilized AIgG (0-300 pg mL-1).

Figure 4. Corrected calibration plot showing changes in electron transfer resistance vs PSA concentration: (a) entrapped APSA 0-300 ng mL-1 PSA, (b) 0-10 ng mL-1 PSA, (c) affinity immobilized APSA 0-300 pg mL-1 PSA, and (d) 0-10 pg mL-1 PSA.

ng mL-1. Upon comparison of Figure 3a,b with Figure 3c,d, it is apparent that approximately 20% of the APSA-affinity immobilized sensor response encountered is, in fact, nonspecific and in the case of the immunosensor with entrapped APSA, 30-40% of the response is nonspecific. 6202

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To obtain a “corrected” calibration profile, in terms of electron transfer resistance change for an APSA immunosensor, nonspecific responses were subtracted from the specific responses over the entire analytical concentration range. In total, 10 APSA immobilized sensors of each type and 10 corresponding AIgG-

immobilized sensors were interrogated over their active concentration range, allowing an assessment of the reproducibility of the responses for this system. Results are presented in Figure 4a-d. The error bars are the standard deviations obtained for the 10 matched sensor pairs from the mean “corrected” values. From Figure 4a, a linear response was observed for the entrapped immunosensors from 1-100 ng mL-1 PSA. Figure 4b shows an expanded view from 1 to 10 ng mL-1 PSA and shows linear behavior with R2 ) 0.998. Affinity immobilized immunosensors displayed a linear range from 1 to 100 pg mL-1 PSA (figure 4c) with R2 ) 0.997 between 1 and 10 pg mL-1 PSA (Figure 4d). Above the linear ranges, the curve tends toward a plateau as the antibody loading becomes saturated. The reproducibility of the “corrected” response obtained from the standard deviation for 10 matched sensor pairings enables discrimination at 1 ng mL-1 (entrapped) or 5 pg mL-1 (affinity) PSA. The limit of detection of the APSA immunosensor (3 times the standard deviation of the baseline value) is 1 ng mL-1 (entrapped) or 1 pg mL-1 (affinity).The much lower limit of detection of sensors produced by affinity immobilization is striking in comparison to those produced via entrapment. The obvious explanation is that affinity immobilization leads to a structure where the antibodies are all at the surface of the polyaniline protrusions and potentially available for binding. Conversely the entrapped antibodies are located at random throughout the “mushroom”, and since both antibody and antigen are large molecules and probably cannot diffuse within the polyaniline matrix, only those located at the microelectrode surface are available for binding. There is the potential for another factor to lower the sensitivity of the entrapped sensors, namely, that deposition must be carried out at a pH of 4.0 to maintain polyaniline in its conductive state and there may be as a result some denaturing of the antibody; however, later results with acid washing of affinity bound antibodies do indicate that the antibodies display some acid stability. Regeneration Studies. Attempts were made so see whether it was possible to regenerate the sensors by acid washing of five APSA-immobilized sensors and five matched nonspecific AIgGimmobilized sensors as detailed within the Supporting Information.To split the antibody-antigen complex and regenerate the sensors, samples were dipped 3 times into 0.1 M HCl/1 min and rinsed with PBS between each dip. A final rinse in 50 mL of PBS produced PSA-free sensors which could then be used for fresh analysis. The results in Figure 5a showing the change in electron transfer resistance following exposure of the affinity sensor to 100 pg mL-1 PSA indicate that negligible changes in “corrected” electron transfer response are encountered after the first regeneration for the five sensor pairings. Both specific and nonspecific sensors exhibited reproducible behavior. After the second regeneration it is apparent that the “corrected” electron transfer change has been significantly lowered with a response approximately 85% that of the initial response. Subsequent regeneration attempts showed significant loss of sensor response. The initial reversibility after one regeneration attempt can be attributed to the restoration of the availability of the binding sites of the antibody by acidic washing and indicates that the avidin-biotin interactions are not disrupted. Unfortunately, a second washing of the sensing layer substantially lowered the

activity of the antibody and subsequent attempts to regenerate the surface removed all analyte recognition properties. Similar experiments were performed using entrapped APSA sensors; however, the first acid wash was found to remove all activity and the sensor proved impossible to regenerate. It is possible in this case that acid washing of the sensors leaches out the antibody (especially if only antibodies near the surface of the microelectrode are active). These results are significant since they demonstrate the potential for the development of the regeneration process for the affinity immobilized sensors. Further work could yield an optimum protocol for further, or perhaps continual, true reversibility. Greater initial response recovery over a larger number of cycles would clearly be beneficial. The phenomenon of reversible regeneration at immunosensor surfaces is a behavior that is rarely encountered and is a highly desirable property of these APSA immunosensors. Stability of the Immunosensors. Stability of the APSA immunosensors was assessed from storage in the dry state, at 4 °C, for a period of 12 weeks. Initially 18 matched sensor pairings of APSA (specific) and AIgG (nonspecific) immunosensors fabricated by both methods were refrigerated at 4 °C in sealed, sterilized vessels. Every 2 weeks a set of 3 matched sensor pairings were removed from storage and interrogated over the full PSA concentration range from 1 to 300 pg mL-1. Figure 5b shows the mean “corrected” electron transfer resistances for the affinity based sensor obtained at 100 pg mL-1 PSA for each time interval. Only minimal changes occur, indicating that the activity of the antibody receptor biocomponent is unaffected by storage over a period as long as 12 weeks. The results lend themselves well to possible packaging and suggests a satisfactory shelf life. Figure 5c shows the results for the entrapped APSA immunosensor. Negligible changes in response (PSA concentration of 100 ng mL-1) are seen up to 6 weeks, after which there is a general decrease in response until after 12 weeks the response is approximately one-third of a fresh sample, indicating a loss of activity. Effect of Interferences. The experimental protocol for the assembly of affinity immobilized immunosensors and testing of PSA was repeated in an identical manner except that the antigen solutions contained one of either 10 ng mL-1 BSA, HSA, CA125, NSE, or S-100[β] along with the PSA. In these cases the calibration results did not differ significantly from the original calibration figures (Figure 4a,b). The mean “corrected” electron transfer changes showed a variation of less than 5% for all concentrations and standard deviations that were comparable to the original. Thus, adsorption of BSA, HSA, CA125, NSE, and S-100[β] and any nonspecific interactions at the sensor surface do not occur or do not affect significantly the biorecognition events (stable immunocomplex formation) and their impedance response. The matched AIgG sensors (nonspecific) also display similar responses to sensors measured without interferences. It is postulated that the HSA blocking during fabrication minimizes nonspecific effects almost totally. The AIgG sensor response is thus not due to the presence of the target analyte or the presence of foreign proteins. This control sensor must be responding to other effects that may be occurring such as changes in the polymer or effects from the Analytical Chemistry, Vol. 80, No. 16, August 15, 2008

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Figure 5. Effect of (a) regeneration and (b) storage on the mean change in electron transfer resistance at 100 pg mL-1 PSA for affinityimmobilized immunosensors, (c) effect of storage on the mean change in electron-transfer resistance at 100 ng mL-1 PSA for entrapped APSA immunosensors.

solution itself on the sensor. These effects are always minor compared to the signal obtained when a sensor immobilized with a specific receptor antibody is exposed to its target analyte. These 6204

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results are promising for prototype development since they suggest that sensing of a target antigen in “real” samples is a plausible proposal.

CONCLUSIONS Our work verifies the fabrication techniques and experimental interrogation protocol as a viable approach toward the labeless sensing of PSA and demonstrates the generic nature of the technology. The APSA-PSA immunosensor system has low nonspecific responses associated with the faradaic sensing of PSA. The affinity-based immunosensor gave a “corrected” linear-type response from 1 to 100 pg mL-1 with saturation occurring above this concentration. The limit of detection of the affinity immunosensor was found to be 1 pg mL-1 with an ability to discriminate at 5 pg mL-1 PSA. The immunosensors formulated by an entrapment technique were much less sensitive (“corrected” linear-type response from 1 to 200 ng mL-1, limit of detection of 1 ng mL-1, and an ability to discriminate at 5 ng mL-1). The affinity based immunosensor was reversible for one regeneration treatment; further treatments have detrimental effects on the sensor performance. Stability studies of the sensors were encouraging with negligible loss of response after storage for 12 weeks at 4 °C in the dry state. No interference was observed upon exposure to high levels of other proteins. The entrapped APSA sensors could not be regenerated and had inferior storage properties (up to 6 weeks). The low quantification levels and limits of detection for PSA have illustrated the potential use of the affinity based microelectrode array immunosensor assemblies for the labeless detection

of PSA. Sensors fabricated using an entrapment method, although less sensitive, were found to be capable of determining PSA levels within the range of clinical significance. Although this initial study shows great promise there still remains further work to be performed. Measurements need to be performed on clinical samples. However it should be noted the extremely low limits of detection of the affinity based sensors indicates that clinical samples can be highly diluted in PBS before use, thereby diluting and minimizing the effects of potential interferences. The minimal effect of high levels of interfering proteins also bodes well for potential applications. ACKNOWLEDGMENT This work, including funding for A.C.B., has been supported by the European Community Grants QLRT-2001-02583 (SMILE) and NMP2-CT-2003-505485, (ELISHA) Framework VI contracts. SUPPORTING INFORMATION AVAILABLE Additional information as noted in the text. This material is available free of charge via the Internet at http://pubs.acs.org.

Received for review March 8, 2008. Accepted June 12, 2008. AC800491M

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