We report a strategy enabling ultrasensitive colorimetric detection of 17β-estradiol (E2) in water and urine samples using DNA aptamer-coated gold nanoparticles (AuNPs). Starting from an established sensor format where aggregation is triggered when target-bound aptamers dissociate from AuNP surfaces, we demonstrated that step-change improvements are easily accessible through deletion of excess flanking nucleotides from aptamer sequences. After evaluating the lowest energy two-dimensional configuration of the previously isolated E2 binding 75-mer aptamer (KD ∼25 nM), new 35-mer and 22-mer aptamers were generated with KD's of 14 and 11 nM by simply removing flanking nucleotides on either side of the inner core. The shorter aptamers were found to improve discrimination against other steroidal molecules and to improve colorimetric sensitivity for E2 detection by 25-fold compared with the 75-mer to 200 pM. In comparing the response of all sequences, we find that the excess flanking nucleotides suppress signal transduction by causing target-bound aptamers to remain adhered to AuNPs, which we confirm via surface sensitive electrochemical measurements. However, comparison between the 22-mer and 35-mer systems show that retaining a small number of excess bases is optimal. The performance advances we achieved by specifically considering the signal transduction mechanism ultimately resulted in facile detection of E2 in urine, as well as enabling environmental detection of E2 at levels approaching biological relevance.

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Despite a large number of publications describing biosensors based on electrochemical impedance spectroscopy (EIS), little attention has been paid to the stability and reproducibility issues of the sensor interfaces. In this work, the stability and reproducibility of faradaic EIS analyses on the aptamer/mercaptohexanol (MCH) self-assembled monolayer (SAM)-functionalized gold surfaces in ferri- and ferrocyanide solution were systematically evaluated prior to and after the aptamer-probe DNA hybridization. It is shown that the EIS data exhibited significant drift, and this significantly affected the reproducibility of the EIS signal of the hybridization. As a result, no significant difference between the charge transfer resistance (R CT) changes induced by the aptamer-target DNA hybridization and that caused by the drift could be identified. A conditioning of the electrode in the measurement solution for more than 12 h was required to reach a stable R CT baseline prior to the aptamer-probe DNA hybridization. The monitored drift in R CT and double layer capacitance during the conditioning suggests that the MCH SAM on the gold surface reorganized to a thinner but more closely packed layer. We also observed that the hot binding buffer used in the following aptamer-probe DNA hybridization process could induce additional MCH and aptamer reorganization, and thus further drift in R CT. As a result, the R CT change caused by the aptamer-probe DNA hybridization was less than that caused by the hot binding buffer (blank control experiment). Therefore, it is suggested that the use of high temperature in the EIS measurement should be carefully evaluated or avoided. This work provides practical guidelines for the EIS measurements. Moreover, because SAM-functionalized gold electrodes are widely used in biosensors, for example, DNA sensors, an improved understanding of the origin of the observed drift is very important for the development of well-functioning and reproducible biosensors.

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