A Multiplexed, Two-Electrode Platform for Biosensing based on
DNA-Mediated Charge Transport
Ariel L. Furst
1
,
Michael G. Hill
1,2
, and
Jacqueline K. Barton
1,*
1
Division of Chemistry and Chemical Engineering, California Institute of Technology, Pasadena,
CA 91125
2
Department of Chemistry and Chemical Biology, Occidental College, Los Angeles, CA 90041
Abstract
We have developed a thin layer, multiplexed biosensing platform that features two working-
electrode arrays for detecting small molecules, nucleic acid sequences, and DNA-binding proteins.
DNA duplexes are patterned onto the primary electrode array, while a secondary electrode array is
used both to initiate DNA monolayer formation, and for electrochemical readout
via
DNA-
mediated charge transport (DNA CT) chemistry. Electrochemical reduction of Cu(phendione)
2
2+
(phendione is 1,10-phenanthroline-5,6-dione) at the secondary electrodes induces covalent
attachment
via
click chemistry of ethynyl-labeled DNA probe duplexes onto the primary
electrodes that have been treated with azide-terminated alkythiols. Electrochemical impedance
spectroscopy and cyclic voltammetry confirm that catalyst activation at the secondary electrode is
essential to maintain the integrity of the DNA monolayer. Electrochemical readout of DNA CT
processes that occur at the primary electrode is accomplished at the secondary electrode. The two-
electrode system enables the platform to function as a collector-generator using either
ferrocyanide or ferricyanide as mediators with methylene blue and DNA charge transport.
Electrochemical measurements at the secondary electrode eliminate the need for large background
corrections. The resulting sensitivity of this platform enables the reliable and simultaneous
detection of femtomoles of the transcription factors TATA-binding protein and CopG on a single
multiplexed device.
INTRODUCTION
Analytical methods for reliable biomolecule detection are becoming increasingly important
with the continued discovery of disease-related biomarkers. Electrochemical nucleic acid-
based assays,
1
–
10
particularly those that utilize DNA-mediated charge transport (DNA
CT)
11
,
12
are especially promising for sensing platforms. Devices based on DNA CT
effectively report on the integrity of the
π
-stacked DNA bases; perturbations to the proper
stacking, resulting from lesions, single nucleotide polymorphisms, or protein binding events
that affect the base stack, attenuate the electrochemical signal. DNA CT has been employed
*
To whom correspondence should be addressed: jkbarton@caltech.edu.
Supporting Information Available: Additional experimental details as well as a CV of the copper precatalyst, impedance spectrometry
data of the DNA monolayers, and spacer height optimization data are available.
HHS Public Access
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. 2015 June 16; 31(23): 6554–6562. doi:10.1021/acs.langmuir.5b00829.
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successfully in the detection of a variety of biomolecules, including DNA fragments,
chemically modified DNA, and DNA-binding proteins, many of which are not specifically
detectable using alternative sensing platforms.
11
,
13
Substrates for DNA-based electrochemical systems typically are prepared by self-
assembling thiolated DNA duplexes onto gold electrodes, followed by backfilling with an
alkylthiol to passivate any remaining surface-exposed gold.
14
–
16
One major challenge with
this methodology is the limited control over monolayer composition, both in the total
amount of DNA assembled and its dispersion within the monolayer.
17
–
19
For biosensing
applications, which rely on direct interactions between target biomolecules and the DNA
duplexes attached to the electrode surface, adequate spacing between the DNA duplexes is
critical to provide the target sufficient access. Inevitably, clustering occurs with thiolated
DNA. Recently, we demonstrated the utility of applying mixed alkylthiol monolayers doped
with variable amounts of azide-terminated functional groups to gain more control over
monolayer formation.
20
Subsequent coupling
via
conjugation of cyclooctyne-labeled DNA
yields surfaces containing evenly dispersed DNA with coverages that mirror the mole
fraction of azide in the underlying film. The resulting monolayers allow greater access of
DNA-binding proteins to individual helices within the films, permitting devices with greater
sensitivity to these biomolecules.
While tethering DNA to surfaces with cyclooctyne provides a strong foundation for more
controlled monolayer formation, ideally, DNA probe molecules would feature a simple
terminal alkyne group to avoid additional synthetic steps. The well known Huisgen 1,3-
dipolar cycloaddition (“click” reaction)
21
catalyzed by copper(I) has been used previously to
form homogenous monolayers using terminal alkyne-labeled probe molecules.
22
,
23
Indeed,
because of the instability of copper(I) in aqueous solution and its reactivity with DNA,
24
–
26
electrochemical methods to generate copper(I)
in situ
from copper(II) precursors have been
developed, and the coupling of alkyne-labeled oligonucleotides to azide-terminated surfaces
via
electrochemically induced click chemistry has been reported.
27
–
30
We have now employed a two-electrode platform in which simple alkyne-labeled duplexes
are coupled to azide-terminated surfaces by copper(I) species generated
in situ
at a
secondary working electrode positioned over the alkylthiol monolayer.
31
Our attempts to
fabricate surfaces suitable for DNA CT using published methods with a single working
electrode were unsuccessful: reliable electrochemical readout of DNA-mediated chemistry
was hampered by interference from the irreversible products of copper(II) reduction at the
modified electrochemical surface (likely adsorption of copper films onto the electrode
surface). Our new method
31
allows for the attachment of multiple DNA sequences onto a
single electrode, with tight control over the probe-molecule spacing.
32
A multiplexed
version of this methodology has enabled the sensitive detection of DNA methyltransferase
activity directly from human tissue samples.
33
Here we report the full characterization of this multiplexed, two working-electrode platform.
In addition to minimizing undesirable copper byproducts at DNA-modified surfaces, we
have found that readout of electrocatalytically generated reporter molecules at the secondary
electrode greatly enhances the sensitivity and specificity of DNA CT assays. Notably, this
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mode of detection eliminates large background signals, and hence the requirement for
sophisticated data processing. This platform enables detection of single base mismatches as
well as the selective and specific detection of two transcription factors, TATA binding
protein (TBP) and CopG, with sensitivities significantly greater than those achieved using
single working-electrode platforms.
EXPERIMENTAL
Design of Experimental Platform
The multiplexed, two-electrode array was fabricated from two ¼” Teflon blocks separated
by a Teflon gasket of various thicknesses (Figure 1). Gold wires (1-mm diameter) were then
inserted into holes drilled into the Teflon to form complementary 5 × 3 electrode arrays on
each block. Each pair of complementary electrodes was 5 mm from its nearest neighbor,
providing the opportunity to isolate each pair into individual wells. The electrodes were
sealed into the Teflon using superglue. The top array featured additional holes (1.5-mm
diameter) to provide the reference and auxiliary electrodes access to the working solution.
Spacers that separated the individual wells were constructed from 1.5 mm thick Teflon.
DNA Attachment to Alkanethiol Monolayers
A 10 mM aqueous Cu(phendione)
2
2+
(phendione=1,10-phenanthroline-5,6-dione) solution
was prepared by combining one equivalent of CuSO
4
(10 μmol, 15.9 mg) with two
equivalents of phendione (20 μmol, 42.0 mg) in 10 mL of deionized H
2
O. ESI-MS: 580.2
(calc: 580.0). The complex was additionally isolated as the PF
6
−
salt. Prior to application to
the electrode surface, the complex was diluted to a final concentration of 1 mM in Tris
buffer (10 mM Tris, 100 mM KCl, 2.5 mM MgCl
2
, 1 mM CaCl
2
, pH 7.6). The catalyst
solution was combined with 5
′
-labeled ethynyl DNA (final concentration of 25 μM), and a
constant potential of −350 mV v. AgCl/Ag was applied to the sensing (top) electrode array
to reduce the Cu(II) and initiate the coupling of the DNA to the azide-terminated
monolayers. The potential was applied for 15 minutes. Multiple sequences of DNA were
attached to the array through the sequential activation of different secondary electrodes. For
example, well matched and mismatched DNA were attached to the same array through the
preliminary activation of secondary electrodes 1 – 9 in the presence of well matched DNA,
followed by rinsing of the platform and subsequent activation of secondary electrodes 10 –
15 in the presence of DNA containing a single-base mismatch.
Characterization of DNA-modified Monolayers
All electrochemical experiments were performed on a CH Instruments 760E bipotentiostat.
For electrochemical impedance spectroscopy experiments, 400 μM potassium ferricyanide
in phosphate buffer (5 mM phosphate, 50 mM NaCl, pH 7.0) was used. For all other
experiments, electrochemistry was conducted in Tris buffer (10 mM Tris, 100 mM KCl, 2.5
mM MgCl
2
, 1 mM CaCl
2
, pH 7.6) with 4 μM methylene blue and 300 μM potassium
ferricyanide; for experiments with covalent Nile Blue, methylene blue was omitted. For
mismatch discrimination and protein binding experiments, constant potential amperometry
was used, with potential applied for 90 s. The primary electrode was held at −400 mV v.
AgCl/Ag, and the secondary electrode was held at 350 mV v. AgCl/Ag.
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RESULTS AND DISCUSSION
We have designed and fabricated an addressable, multiplexed biosensing array that features
two sets of complementary electrodes separated by a thin film (Figure 1). The bottom
electrodes, which comprise the primary array, are modified with covalently bound DNA
sequences dispersed within a mixed alkylthiol monolayer, while the top electrodes, which
form the secondary array, are unmodified and are used both for activating a DNA-coupling
catalyst and for the electrochemical readout. An overview of the steps required for the
primary electrode-array modification, DNA coupling, and electrocatalytic detection using
this two-electrode sensing platform are summarized in Scheme 1.
Electrochemical Response of the Coupling Catalyst
The electrochemical initiation of azide/alkyne coupling
via
copper(II) reduction has been
explored previously.
34
While several chelating ligands for copper(I) click chemistry have
been reported, our experiments have focused on the bipyridyl derivative 1,10-
phenanthroline-5,6-dione (phendione), as this ligand is commercially available and yields a
water-soluble complex. The cyclic voltammogram (CV) of [Cu(phendione)
2
][SO
4
] in Tris
buffer shows several copper-centered reductions between + 0.20 and −0.40 V
vs.
AgCl/Ag
(Figure S1). Importantly, these processes are only partially chemically reversible. The CV
also shows a coupled oxidative response with a shape characteristic of anodic stripping,
suggesting that the electrochemical reactions result in the adsorption of at least some copper-
containing species onto the electrode surface. The deposition of copper following
electrochemical reduction is further supported by the formation of a visible, black surface
film following the application of potentials in the range of −0.30 to −0.40 V
vs.
AgCl/Ag.
Formation of DNA Monolayers with Activation from Primary vs. Secondary Electrodes
DNA-modified surfaces were prepared on the primary array using a two-step process that
involved: (i) the self-assembly of a mixed alkylthiol monolayer containing 50% azide and
50% phosphate head groups onto the primary electrodes, followed by (ii) the
electrochemical reduction of Cu(phendione)
2
2+
at the
top
(secondary) electrodes in the
presence of alkyne-labeled DNA duplexes. Electrogenerated copper(I) induces covalent
attachment of the DNA
via
alkyne/azide coupling (Scheme 1). Based on previous studies, an
underlying 1:1 ratio of azide-to-phosphate head groups provides adequate spacing between
the individual helices for substrate access within the DNA monolayer, while still
maintaining a sufficient concentration of DNA on the surface for reliable detection.
20
Importantly, before attachment of the DNA duplexes, the azide/phosphate-modified primary
electrodes showed no electrochemical response in the presence of up to 8 μM methylene
blue (MB) and/or 500 μM Fe(CN)
6
3−
.
The catalytic precursor, Cu(phendione)
2
2+
(1 mM), was activated by applying a potential of
−350 mV (
vs
. AgCl/Ag) to the secondary electrodes for 15 minutes. After rinsing the system
with Tris buffer, cyclic voltammetry at the primary electrodes revealed reproducible and
featureless background currents, characteristic of DNA-modified electrodes. No signals
attributable to copper-reduction byproducts were observed. By contrast, if the
Cu(phendione)
2
2+
activation was carried out at the
primary
(bottom) electrodes instead, the
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resulting cyclic voltammograms exhibited large, sloping, and irreproducibly “bumpy”
signals, even after multiple rinsings. This observation suggests partial (and variable)
passivation of the primary-electrode surfaces due to the irreversible copper(II)
electrochemistry.
Electrochemical impedance spectroscopy (EIS)
36
was also performed to evaluate the effect
of copper-film deposition on the electrochemical properties of DNA-modified electrodes.
Data were collected at DNA-modified surfaces prepared
via
Cu(phendione)
2
2+
activation at
both the primary and secondary electrodes, and compared to analogous data gathered at bare
gold and at electrodes modified with only the underlying mixed alkylthiol monolayers. EIS
was conducted at the Fe(CN)
6
3−/4−
redox couple, following literature protocols for the label-
free detection of DNA.
35
,
36
Nyquist plots (Figure S2) constructed from measurements
recorded at both bare gold surfaces and electrodes modified only with the underlying
alkylthiols display a small impedance arc, consistent with low surface capacitance and a
response dominated by the diffusion of Fe(CN)
6
3−/4−
species to the surface.
36
–
39
Alternatively, the conjugation of DNA duplexes onto the primary gold monolayers provides
an electrostatic barrier that effectively passivates the negatively charged Fe(CN)
6
3−/4−
complexes from the electrode surface; the corresponding Nyquist plots therefore show a
significantly larger capacitive arc. Notably, the plots from DNA monolayers formed
via
catalyst activation at the primary electrodes show the largest (and least reproducible) arcs,
indicating higher electron-transfer resistance through these films than through the analogous
DNA monolayers prepared
via
activation from the secondary array (Figure S2). These
results are fully consistent with the cyclic voltammetry of Cu(phendione)
2
2+
, and with the
visible appearance of black deposits on whichever array was used to activate the copper(II).
Importantly, when the secondary electrodes were used for activation, the precipitate could
be easily removed by polishing the array prior to detection, without perturbing the DNA
monolayers on the primary-electrode surfaces.
Electrochemical Readout at the Primary Electrode
In order to amplify the electrochemical readout from DNA CT-based sensors, we previously
developed an electrocatalytic cycle using MB and Fe(CN)
6
3−
in solution.
12
The cycle begins
with the DNA-mediated electrochemical reduction of intercalated MB. The product of that
reaction, leucomethylene blue (LB), has a lower affinity for DNA and dissociates from the
π
-stack. LB then reduces Fe(CN)
6
3−
to Fe(CN)
6
4−
in solution, thereby regenerating MB that
re-intercalates into the film to begin the cycle again. Under certain experimental
conditions,
39
the measured current is limited only by the diffusion of ferricyanide in
solution, providing dramatically larger signals than those obtained by the direct surface
electrochemistry of MB at the DNA-modified electrode.
In order to evaluate the utility of this electrocatalytic cycle in our two-electrode platform, we
prepared a series of primary electrodes that contained either well-matched (WM) or
mismatched (MM) 18-base-pair DNA duplexes (see SI for sequences). We then monitored
the thin-film MB/Fe(CN)
6
3−
electrochemical response at both the primary and secondary
electrode arrays. As DNA CT is sharply attenuated by lesions within the
π
-stack, the ability
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to discriminate between WM and MM monolayers provides strong evidence for a DNA-
mediated process.
Figure 2 shows representative results when the current is measured at the primary electrode.
Because of the low surface coverage of DNA, the relatively large background currents, and
the restricted diffusion of Fe(CN)
6
3−
owing to the proximity of the secondary electrode, the
cyclic voltammograms do not show the characteristic electrocatalytic wave shapes observed
in the absence of the secondary array. Nevertheless, integrating the measured current traces
reveals a marginally higher net-charge passed through the well-
vs.
mismatched-DNA films.
(As also illustrated in Figure 2, if instead the DNA films are formed
via
copper(II) activation
at the primary electrodes, the background signals are even larger, and the differences in
charge passed through well-
vs.
mismatched DNA are correspondingly smaller.) Thus, while
mismatch discrimination can be accomplished using electrochemical readout at the primary
electrodes, the signal amplification afforded by the addition of Fe(CN)
6
3−
is essentially only
stoichiometric, so the MB/Fe(CN)
6
3−
signals remain very small relative to the much larger
background currents. As with other systems where a single electrode is used both for target
capture
and
electrochemical readout, reliable mismatch detection at the primary array would
require extensive background subtraction of the large non-Faradaic currents that dominate
the electrochemical response in order to discern the subtle differences in the (much smaller)
Faradaic signals that occur at well-
vs.
mismatched sequences.
Electrochemical Readout at the Secondary Electrode
In contrast, using a secondary array for electrochemical readout offers an inherently more
sensitive detection platform, with virtually no interference from background currents.
Because the secondary-electrode array effectually traps solution-bourn analytes within the
thin layer created between the primary and secondary electrodes, holding the secondary
electrodes at potentials positive of the Fe(CN)
6
3−/4−
couple should continually replenish
Fe(CN)
6
3−
in the vicinity of the DNA films by re-oxidizing Fe(CN)
6
4−
as soon as it is
generated during the electrocatalytic cycle. This would enable multiple rounds of turnover
and a correspondingly higher level of amplification. Although the rate of MB/Fe(CN)
6
3−
electrocatalysis is limited by the DNA-binding dynamics of MB under the experimental
conditions used in this study,
39
the two-electrode platform offers, in principle, the
possibility of secondary-electrode readout currents that are larger than would be possible
using passive diffusion.
Indeed, as demonstrated in Figure 3, poising the secondary electrode at +0.35 V (
i.e.,
0.1 V
positive of the Fe(CN)
6
3−/4−
couple) during the electrocatalytic assay enables the system to
function in this “collector-generator” mode: upon reduction of Fe(CN)
6
3−
in solution by the
LB produced
via
DNA-mediated reduction of MB at the primary electrode, the resulting
Fe(CN)
6
4−
is immediately re-oxidized at the secondary electrode to generate a large, “turn-
on” readout current. In contrast, if no DNA CT occurs at the primary electrode, no current is
generated at the secondary electrode, negating the need for any background correction.
Because of the thin layer created by sandwiching the primary and secondary arrays together,
the platform can function as a collector-generator when starting either with Fe(CN)
6
4−
(Figure 3b), as was done previously using scanning electrochemical microscopy, or with
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Fe(CN)
6
3−
(Figure 3a), which is rapidly converted to Fe(CN)
6
4−
via
DNA CT at the primary
electrode surface.
Note that for this platform to function as a collector-generator, the distance between the two
sets of electrodes must be sufficiently close to prevent the Fe(CN)
6
3−/4−
species from
diffusing away. In scanning electrochemical microscopy (SECM), the optimal distance
between the tip and the electrode surface is determined by an approach curve.
40
,
41
Here, we
have experimentally determined the optimal distance between our two sets of working
electrodes by carrying out the analysis using various thicknesses of the Teflon spacer that
separates the two arrays (Figure S3). Empirically, the spacer that simultaneously yielded the
largest readout currents at the secondary electrodes while providing the greatest level of
mismatch discrimination was 127 μm thick, which we therefore used for all experiments. It
may be the case that if the spacer is too small, the electrode arrays are too close to one
another, and our signals are no longer DNA-mediated. This is likely the case with the 50 μm
spacer, with which we do not see mismatch discrimination.
Single-Base Mismatch Detection with Non-covalent and Covalent Redox Probes
To evaluate the ability of this platform to detect single-base mismatches, we prepared a
primary-electrode array that featured the side-by-side conjugation of both well- and
mismatched alkyne-labeled DNA duplexes at adjacent electrode sites. DNA addressing was
readily accomplished by activating the copper(II) coupling catalyst only at specific
(complementary) secondary electrodes for each DNA sequence used. As shown in Figure 4,
electrochemical readout was then carried out using constant-potential amperometry at the
secondary electrodes, in the presence of MB/Fe(CN)
6
3−.
33
The differences in readout
currents generated by well-
vs.
mismatched duplexes are both large and robust. (For
comparison, if the monolayers are instead addressed by copper(II) activation at the primary
electrodes, the resulting readout currents are small, and mismatch discrimination is poor;
this further highlights the importance of catalyst activation from the secondary electrode.)
These signal differences obviate the need for dramatic background subtractions, essentially
providing an on/off sensor for mismatch detection.
We also investigated the possibility of replacing MB with a covalent redox probe, which
may be necessary for some biomolecule detection applications. Nile Blue, a covalent
reporter molecule that is electronically conjugated to the DNA
π
-stack,
48
was previously
found to mediate the electrocatalytic reduction of Fe(CN)
6
3−
in solution. As shown in Figure
5, replacing MB with NB in the two-electrode platform still allows for robust mismatch
discrimination, despite the fact that the NB is covalently tethered to DNA. Based on
electrochemical readout at the secondary electrodes, a 60±10% decrease in the current upon
incorporation of a single-base mismatch occurs with the covalent Nile blue redox probe, as
compared to an 80±10% decrease with noncovalent MB. This difference in mismatch
discrimination is likely due to the higher binding stoichiometry of methylene blue to DNA,
as well as its enhanced mobility within the film. It is therefore unsurprising that MB yields
larger signals and a correspondingly larger differential. This result is consistent with
previous single-electrode studies, in which larger differentials are observed with free probes
than with covalent probes.
48
,
49
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Detection of DNA-binding Proteins
To establish the relevance of this platform for biomolecule detection, DNA-binding protein
detection was investigated using two transcription factors, TBP and CopG. Both proteins
bind to specific sequences of DNA, kinking the duplex to a large degree, and are ideal
targets for DNA CT-based assays. The transcription factor TBP (TATA-binding protein, a
subunit of the eukaryotic TFIID transcription factor) was previously employed as a measure
of sensitivity for a DNA CT-based protein-detection platform.
50
,
51
TBP kinks DNA by over
80° when bound to its TATA target sequence,
52
destacking the DNA bases and disrupting
DNA CT.
20
Likewise, the transcription repressor CopG, which binds DNA as a tetramer at
an ACGTxxxxxACGT site, bends the helix up to 120° and also has a nano-molar binding
affinity.
53
–
55
To determine detection limits using the two-electrode assay, each protein was titrated
individually into the two-electrode platform, and electrochemical readout at the secondary-
electrode during MB/Fe(CN)
6
3−
electrocatalysis was used to monitor protein binding. The
titration data, shown in Figure 6, indicate a greater than 30% decrease in the readout signals
associated with both proteins at concentrations as low as 10 nM; based on the small volumes
required to fill the thin layer between the two electrode arrays, this translates to less than 50
femtomoles of protein. From these titration curves, dissociation constants can be calculated
for protein binding at the DNA-modified surfaces, using a cooperative binding model (the
Hill model).
20
TBP has a surface K
D
of 14±2 nM, and CopG has a surface K
D
of 17±4 nM.
Both TBP and CopG are therefore detectable at concentrations near their solution K
D
’s (3.3
nM and 10 nM, respectively).
50
Notably, the multiplexed nature of this sensing platform enables the specific detection of
multiple proteins simultaneously, using a single primary-electrode array. To illustrate this
capability, each transcription factor was added to a primary array that featured: (i) three
electrodes modified with a non-binding DNA sequence; (ii) six electrodes modified with
DNA containing a TBP binding site; and (iii) six electrodes modified with DNA containing
a CopG binding site. The assay itself was carried out by introducing one protein (either TBP
or CopG) first onto the platform, then measuring the secondary-electrode readout at
addresses that correspond to the locations of the DNA-binding sequences, as well as at the
locations of the non-binding sequences (the positive controls). Any signal attenuation was
measured relative to the readout current at the positive-control addresses. The surface was
subsequently rinsed, the second protein was added, and the readout signal attenuation was
again measured. In each case, signal attenuation occurred only at secondary electrodes that
were complementary to the primary addresses which featured the binding sequence of the
protein added. This observation verifies that changes are due only to the specific binding of
each protein. We note that greater variability between protein-detection experiments is
observed in the dual protein detection experiments, as compared to the individual titrations,
likely because of the added normalization of these data to a non-binding positive control
sequence. This additional normalization is necessary for combined protein experiments to
further ensure that signal decreases are due to specific binding of only one of the proteins.
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CONCLUSIONS
We have described a highly sensitive, multiplexed, two-working-electrode platform for
DNA CT-based electrochemical detection. Using
in situ
electrochemical activation of a
copper(II) catalyst for Huisgen 1,3-dipolar coupling, this platform enables the addressing of
low-density DNA monolayers onto specific sites within a single, primary sensing array.
Catalyst activation at a secondary electrode is essential to maintain the integrity of the DNA,
as shown by cyclic voltammetry, EIS, and constant-potential amperometry. Detection and
electrochemical readout from the secondary-electrode array is similarly necessary to provide
high sensitivity without large background signals. Operating in collector-generator mode
using the well-known MB/Fe(CN)
6
3−
electrocatalytic cycle for DNA-modified surfaces, this
platform provides a “signal-on” detection assay for DNA CT, with virtually no background
currents whatsoever. Notably, this platform allows the specific and simultaneous detection
of femtomoles of the transcription factors, TBP and CopG in a single electrochemical assay.
The new multiplexed, two-electrode detection platform significantly broadens the scope and
applications for detection using DNA CT.
Supplementary Material
Refer to Web version on PubMed Central for supplementary material.
Acknowledgments
We are grateful to the NIH (GM61077) for their financial support of this research.
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