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Link to original content: https://pubmed.ncbi.nlm.nih.gov/35074874
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. 2022 Feb 1;119(5):e2112812119.
doi: 10.1073/pnas.2112812119.

Molecular electronics sensors on a scalable semiconductor chip: A platform for single-molecule measurement of binding kinetics and enzyme activity

Carl W Fuller  1 , Pius S Padayatti  1 , Hadi Abderrahim  1 , Lisa Adamiak  1 , Nolan Alagar  1 , Nagaraj Ananthapadmanabhan  1 , Jihye Baek  1 , Sarat Chinni  1 , Chulmin Choi  1 , Kevin J Delaney  1 , Rich Dubielzig  1 , Julie Frkanec  1 , Chris Garcia  1 , Calvin Gardner  1 , Daniel Gebhardt  1 , Tim Geiser  1 , Zachariah Gutierrez  1 , Drew A Hall  1   2 , Andrew P Hodges  1 , Guangyuan Hou  1 , Sonal Jain  1 , Teresa Jones  1 , Raymond Lobaton  1 , Zsolt Majzik  1 , Allen Marte  1 , Prateek Mohan  1 , Paul Mola 2nd  1 , Paul Mudondo  1 , James Mullinix  1 , Thuan Nguyen  1 , Frederick Ollinger  1 , Sarah Orr  1 , Yuxuan Ouyang  1 , Paul Pan  1 , Namseok Park  1 , David Porras  1 , Keshav Prabhu  1 , Cassandra Reese  1 , Travers Ruel  1 , Trevor Sauerbrey  1 , Jaymie R Sawyer  1 , Prem Sinha  1 , Jacky Tu  1 , A G Venkatesh  1 , Sushmitha VijayKumar  1 , Le Zheng  1 , Sungho Jin  3 , James M Tour  4 , George M Church  5 , Paul W Mola  1 , Barry Merriman  6
Affiliations

Molecular electronics sensors on a scalable semiconductor chip: A platform for single-molecule measurement of binding kinetics and enzyme activity

Carl W Fuller et al. Proc Natl Acad Sci U S A. .

Erratum in

Abstract

For nearly 50 years, the vision of using single molecules in circuits has been seen as providing the ultimate miniaturization of electronic chips. An advanced example of such a molecular electronics chip is presented here, with the important distinction that the molecular circuit elements play the role of general-purpose single-molecule sensors. The device consists of a semiconductor chip with a scalable array architecture. Each array element contains a synthetic molecular wire assembled to span nanoelectrodes in a current monitoring circuit. A central conjugation site is used to attach a single probe molecule that defines the target of the sensor. The chip digitizes the resulting picoamp-scale current-versus-time readout from each sensor element of the array at a rate of 1,000 frames per second. This provides detailed electrical signatures of the single-molecule interactions between the probe and targets present in a solution-phase test sample. This platform is used to measure the interaction kinetics of single molecules, without the use of labels, in a massively parallel fashion. To demonstrate broad applicability, examples are shown for probe molecule binding, including DNA oligos, aptamers, antibodies, and antigens, and the activity of enzymes relevant to diagnostics and sequencing, including a CRISPR/Cas enzyme binding a target DNA, and a DNA polymerase enzyme incorporating nucleotides as it copies a DNA template. All of these applications are accomplished with high sensitivity and resolution, on a manufacturable, scalable, all-electronic semiconductor chip device, thereby bringing the power of modern chips to these diverse areas of biosensing.

Keywords: CMOS chip; biosensor; molecular electronics; single-molecule detection; single-molecule sequencing.

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Conflict of interest statement

Competing interest statement: All authors having the Roswell affiliation (affiliation “a”) are employed by Roswell Biotechnologies, San Diego, CA 92121.

Figures

Fig. 1.
Fig. 1.
Molecular electronic sensor and chip concept. (A) Senor concept: Given a pair of molecules that undergo an interaction, one of the pair is selected as a probe molecule and conjugated to a precision molecular wire (here a synthetic α-helical protein) that spans a nanoscale gap between metal nanoelectrodes. These connect it to a driving voltage source and current monitoring circuit to provide real-time readout of current vs. time, for the current passing through the molecular wire/probe complex. When the target molecule binds to the probe, the resistance of the complex changes, resulting in an observed change in current. The resulting current trace has on/off pulses that provide a direct representation of the molecular interactions. (B) CMOS chip device: A large-scale array of sensors are fabricated on the surface of a CMOS chip. Shown is an annotated image of the CMOS chip device used in these studies. This chip has 16,000 sensors and the circuitry needed to digitize and transfer sensor readings off-chip, at a rate of 1,000 frames per second. (C) SEM image of sensor nanoelectrodes, showing the 20 nm gap for the molecular bridge. Nanoelectrodes shown are fabricated by photolithography, using CMOS foundry-compatible process.
Fig. 2.
Fig. 2.
The DNA hybridization binding sensor. (A) A single-stranded DNA (ssDNA) oligo probe, here the17-mer oligonucleotide (5′-TACGTGCAGGTGACAGG-3′), is conjugated to the bridge using conventional click chemistry at the 5′ end. (B) Example current vs. time trace, showing 6 s of data sampled at 1 kHz, taken with the sensor exposed to a 20 nanomolar (nM) concentration of the target oligo, here the complementary 14-mer strand (5′-CCTGTCACCTGCAC), suspended in a standard buffer solution. Each pulse of current above the baseline represents a single DNA binding event. The durations of the events and time between events are stochastic, with exponential distributions as summarized in SI Appendix, Figs. S2–S4. (C) The distribution of measured current values in the trace, shown alongside the trace as a vertical histogram, provides a visualization of the time spent bound (higher current) and unbound (lower current). In this example, the fraction of the time spent bound is 22%. (D) Response of DNA binding sensor to target concentration: binding of the 17-mer ssDNA probe on the sensor to the 14-mer target, at target concentrations of 10 nM, 100 nM, and 1,000 nM. The width of the peaks (dwell time) remains constant (~25 ms), but the time between peaks decreases from 45 to 4 ms as the concentration is increased, reflecting more frequent concentration-driven interactions. The fraction of time in the bound state (labeled “fraction bound” in plots) was estimated from the current measurement value histograms (Insets at Right). (E) Measurements of dwell time and fraction bound vs. concentration. As expected for DNA hybridization binding, dwell time remains constant with concentration, but fraction of time bound shows a classic saturation curve, from which is computed a binding affinity, Kd, of 39 nM. (F) Single-molecule thermal melting curves derived from the DNA binding sensor. In this experiment, a 45-mer ssDNA probe (5′-CGATCAGGCCTTCACAGAGGAAGTATCCTGTCGTTTAGCATACCC-3′) is attached to the bridge at the 5′ end. Two different complementary target oligos that are closely matched in properties but having different melting points were constructed by using a 15-mer target (5′-CCTCTGTGAAGGCCT) of the 45-mer, and an extension of this to a 20-mer target (5′-CCTCTGTGAAGGCCTGATCG). These were added to the chip at a concentration of 20 nM, performed in series. For each solution, the solution temperature was swept, in 2 °C steps, from 41 °C to 55 °C. Standard DNA melting curves were fit and used to derive the empirical melting points, Tmobs, shown. The results agree with the classical bulk predictions for the difference in Tm between the oligos, but here are measured entirely in a single-molecule context. (G) Specificity of the DNA binding sensor for mismatched DNA. The sensor was used to probe mismatched targets, using the 45-mer hybridization probe. The targets were 20-mers having 0, 1, 2, or 3 mismatches as shown (sequence Inset), added sequentially. The result shows a significant downward trend in fraction bound, as the number of mismatched nucleotides is increased.
Fig. 3.
Fig. 3.
A survey of diverse molecular electronic sensors for binding and enzyme activity, shown to scale in molecular renderings, along with corresponding summary experimental results. (A–D) Protein and small-molecule binding kinetics. As a model system for showing protein binding and small-molecule binding, this sensor is configured to observe (A) a DNA polymerase binding a primer/template, and (C) a nucleotide binding into the polymerase pocket. For this model system, a 17-mer ssDNA template is conjugated to the peptide bridge at its 5′ end (and with 3′ end blocked to prevent the polymerase from binding that site). A complementary 14-mer primer strand is then bound to this, on the distal end of the 17-mer, to create a primer site on the sensor with the 3′-OH available for polymerase binding. (B) Summary kinetics (dwell time, fraction of time bound) for Klenow DNA polymerase binding to the primer site, as polymerase concentration is titrated from 0.008 to 3.8 µM, in a background of 100 nM 14-mer primer to suppress primer dissociation. The inferred binding affinity of the polymerase is Kd = 530 nM. (D) A nucleotide titration is performed to observe the binding in the polymerase pocket, in a noncatalytic buffer so as to observe the binding kinetics without incorporation. A 45-mer template is on the bridge, and a 31-mer primer was bound to the distal end of the template, so that the first template base (A) is complementary to the nucleotide being tested (T). The nucleotide was added in concentrations of 2.5, 5, and 15 µM along with the 100 nM polymerase and primer in the presence of a buffer that has 10 mM Sr2+ (without Mg2+), in which nucleotide incorporation cannot occur. In this buffer, the dNTP will repeatedly bind and dissociate from the polymerase pocket, and the resulting summary binding kinetics are shown. This also serves to illustrate the detection of small molecule binding. (E and F) Aptamer sensors: Aptamer sensors were constructed, here targeting the SARS-CoV-2 S protein with a 57-mer DNA aptamer (E) and targeting the SARS-CoV-2 N protein with a 97-mer DNA aptamer (SI Appendix, Fig. S7B-2), both taken from the literature. (F) The concentration response titration curves for both the S aptamer and N aptamer sensors, for a range of applied target protein concentrations. The binding affinities, Kd, derived from these curves (6.4 nM, 39 nM) are similar to those reported for standard bulk aptamer binding assays in solution. (G and H) Antibody–antigen sensors. As a model system, a fluorescein–antifluorescein antigen–antibody pair was used, with the fluorescein antigen presented on the sensor by tethering it to the bridge using a ssDNA oligo as a linker. A 45-mer oligonucleotide was used, with the 3′ (distal) end of the DNA capped with a fluorescein during synthesis. A commercial antifluorescein antibody (Fab) was added in TKS buffer on the chip. The summary kinetics are shown for dwell time and fraction of time bound, as the concentration of antibody is titrated over the range shown. The inferred binding affinity was Kd = 1.3 µM. It was observed that all binding signals were extinguished when 4 mM of free fluorescein was added to saturate the antibody, verifying the specificity of the binding signal. (I and J) A CRISPR/Cas enzyme activity sensor. To assemble a Cas enzyme as a probe on the bridge, first a guide RNA targeting a dsDNA target for a CRISPR/Cas12a enzyme was conjugated to the bridge, and these were assembled on chip. A Cas12a enzyme was provided in solution and allowed to dock to the guide RNAs on the bridges, thereby programming it for the target dsDNA, and also effectively tethering it to the bridge as a probe. For these experiments, the guide RNA is a 40-mer, attached to the bridge peptide using click chemistry at the 13th nucleotide, which is the base that extends furthest outside the enzyme in the pseudoknot loop. The kinetics are summarized in the titration curve, showing fraction of time bound saturating as the dsDNA target varies in concentration, in the presence of a concentration of 20 nM free (untargeted) Cas12a enzyme. Thus, this configuration acts directly as a sensor for the dsDNA target, without assessing posttarget-binding nonspecific single-stranded nuclease activity. This latter nonspecific activity is also observable on the sensor when provided with a ssDNA substrate. The observed binding affinity for the dsDNA target is Kd = 3 pM. The experimental buffer was 20 mM Tris·HCl pH 8.0, 20 mM KCl, 10 mM SrCl2, 4 mM DTT. Two examples of raw sirgnal traces for D (small molecule) and E (protein) are included in SI Appendix, Fig. S7 B and F, showing that the character of the bridge current signals is similar for these diverse probes.
Fig. 4.
Fig. 4.
DNA Polymerase Activity Sensor. A phi29 DNA polymerase is conjugated to the sensor bridge using the SpyTag-SpyCatcher conjugation scheme shown. The 25-second-long signal trace shows an isolated burst of sensor activity that occurs after adding a primed 40-mer template (sequence 5'-25T-15G-3') and corresponding (dCTP and dATP) nucleotides. The expectation is the polymerase would acquire a template and incorporate 15 C's followed by 25 A's. A series of ~40 discrete major pulses are seen, representing putative incorporation events. The signal trace has ~15 wide-spaced, narrower pulses on the left (green region), and ~25 closely spaced, broader pulses on the right (red region), suggesting these are the C and A events, respectively, and that therefore the C and A incorporations events can be distinguished by examining pulse features.
Fig. 5.
Fig. 5.
Using a DNA binding sensor to detect a viral target under mock assay conditions. The DNA binding probe chip is used to detect a target DNA PCR product produced from a contrived sample, with confounding complex background. (A) The titration curves show chip sensor response, for various concentrations of targets, for targets consisting of a synthetic 24-mer positive control, an unpurified PCR product from a contrived saliva sample, and this PCR product with salmon sperm DNA added at 2 μg/mL, to mimic the impact of having background genomic DNA contamination in saliva. This high, complex background had little impact on the results. (B) A 3 s signal trace at the lowest concentration tested, 100 pM, showing 5.2% fraction of time bound. This illustrates the strong signal spikes, and the potential to detect much lower concentrations through longer observations.
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