Folding-based electrochemical biosensors: the case for responsive nucleic acid architectures

Abstract

Biomolecular recognition is versatile, specific, and high affinity, qualities that have motivated decades of research aimed at adapting biomolecules into a general platform for molecular sensing. Despite significant effort, however, so-called "biosensors" have almost entirely failed to achieve their potential as reagentless, real-time analytical devices; the only quantitative, reagentless biosensor to achieve commercial success so far is the home glucose monitor, employed by millions of diabetics. The fundamental stumbling block that has precluded more widespread success of biosensors is the failure of most biomolecules to produce an easily measured signal upon target binding. Antibodies, for example, do not change their shape or dynamics when they bind their recognition partners, nor do they emit light or electrons upon binding. It has thus proven difficult to transduce biomolecular binding events into a measurable output signal, particularly one that is not readily spoofed by the binding of any of the many potentially interfering species in typical biological samples. Analytical approaches based on biomolecular recognition are therefore mostly cumbersome, multistep processes relying on analyte separation and isolation (such as Western blots, ELISA, and other immunochemical methods); these techniques have proven enormously useful, but are limited almost exclusively to laboratory settings. In this Account, we describe how we have refined a potentially general solution to the problem of signal detection in biosensors, one that is based on the binding-induced "folding" of electrode-bound DNA probes. That is, we have developed a broad new class of biosensors that employ electrochemistry to monitor binding-induced changes in the rigidity of a redox-tagged probe DNA that has been site-specifically attached to an interrogating electrode. These folding-based sensors, which have been generalized to a wide range of specific protein, nucleic acid, and small-molecule targets, are rapid (responding in seconds to minutes), sensitive (detecting sub-picomolar to micromolar concentrations), and reagentless. They are also greater than 99% reusable, are supported on micrometer-scale electrodes, and are readily fabricated into densely packed sensor arrays. Finally, and critically, their signaling is linked to a binding-specific change in the physics of the probe DNA, and not simply to adsorption of the target onto the sensor head. Accordingly, they are selective enough to be employed directly in blood, crude soil extracts, cell lysates, and other grossly contaminated clinical and environmental samples. Indeed, we have recently demonstrated the ability to quantitatively monitor a specific small molecule in real-time directly in microliters of flowing, unmodified blood serum. Because of their sensitivity, substantial background suppression, and operational convenience, these folding-based biosensors appear potentially well suited for electronic, on-chip applications in pathogen detection, proteomics, metabolomics, and drug discovery.

Figure 1

The first E-DNA sensor comprised a redox-tagged stem-loop DNA attached to an interrogating electrode. In the absence of target, the redox tag is held in proximity to the electrode, ensuring efficient electron transfer (eT) and a large, readily detectable faradaic current. Upon hybridization with a target the redox tag is removed from the electrode, impeding the signaling current.

Figure 2

E-DNA sensors are rapid, selective and readily reusable. (Top) For example, an E-DNA sensor specifically responds to target even in the presence of 50,000× excess genomic DNA and when deployed directly in complex “dirty” samples such as clinical materials, soil-suspensions and foodstuffs. (Bottom) E-DNA sensors equilibrate within minutes and are regenerated via a simple, 30 s wash with deionized water.

Figure 3

The signal gain of E-DNA sensors is a function of the density with which the probe DNAs are packed onto the electrode surface. Specifically, signaling improves with increasing probe density, presumably because crowding between neighboring probe-target duplexes minimizes electron transfer from the bound state, resulting in increased signal change upon hybridization.

Figure 4

This “signal-on” E-DNA architecture, based on a target-induced strand displacement mechanism, achieves sub-picomolar detection limits. In this mechanism sensor target binding displaces a flexible, single-stranded element modified with a redox-tag. This, in turn, strikes the electrode, generating a large increase in faradaic current.

Figure 5

The first electrochemical, aptamer-based (E-AB) sensor comprised a redox-tagged DNA aptamer directed against the blood-clotting enzyme thrombin. Thrombin biding reduces the current from the redox tag, readily signaling the presence of the target.

Figure 6

A signal-on E-AB sensor based on the strand-displacement mechanism and directed against the protein thrombin achieves a 10-fold increase in signal gain over its signal-off counterpart (Fig. 5) and, in turn, a 7-fold increase in sensitivity.

Figure 7

Small molecule E-AB sensors support real-time detection even in complex sample matrices, such as blood serum. The cocaine E-AB sensor, for example, supports real-time detection of cocaine directly in undiluted blood serum as it flows through a sub-microliter chamber. The letter designations on the external leads denote the “reference,” “counter,” and “working” electrodes to which they are attached.

Figure 8

Utilizing double-stranded DNA as a support scaffold for a small molecule receptor, we have fabricated E-DNA-like sensors for the detection of protein-small molecule interactions. Shown, for example, is the detection of, low nanomolar concentrations of antibodies against the drug digoxigenin.