Label-free polypeptide-based enzyme detection using a graphene-nanoparticle hybrid sensor

Article abstract of DOI:10.1002/adma.201202961

A graphene-nanoparticle (NP) hybrid biosensor that utilizes an electrical hysteresis change to detect the enzymatic activity and concentration of Carboxypeptidase B was developed. The results indicate that the novel graphene-NP hybrid biosensor, utilizing

Full text of DOI:10.1002/adma.201202961

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Label-Free Polypeptide-Based Enzyme Detection
Using a Graphene-Nanoparticle Hybrid Sensor
Sung Myung, Perry T. Yin, Cheoljin Kim, Jaesung Park, Aniruddh Solanki,
Pavel Ivanoff Reyes, Yicheng Lu, Kwang S. Kim, and Ki-Bum Lee*
Over the last few decades, the carbon nanomaterial-based elec-
trical detection of chemical and biological molecules has gained
much attention with its extensive applications to genomic,
a zero-band gap. It forms ideal two-dimensional carbon crystals
that show high electrical conductivity in the absence of any
doping even at room temperature. Graphene also demonstrates
superb mechanical flexibility and high chemical/thermal sta-
bility. Furthermore, due to its low noise level, graphene is
extremely sensitive to perturbations in electrical conductance
that are induced by biological reactions. These biological reac-
tions can include, but are not limited to, the enzyme-mediated
degradation of a target substrate and the binding of proteins
to other biomolecules. There have been a number of graphene-
proteomic and environmental analysis as well as for clinical
diagnosis.[
1–12]
Typical carbon nanomaterial-based electrical
detection systems are composed of two components: a bio-
logical recognition element and a signal transduction element.
These devices can detect minute concentrations of target mol-
ecules by measuring the change in electrical conductance that
results from the binding of target molecules to the surface of
the carbon nanomaterial.[
13–17]
However, the sensitivity as well
based protein sensors with detection limits on the μM scale.
[15–17]
as selectivity of these devices depends heavily on the electrical
properties of the nanomaterial used in the system and on the
interfacial chemistry that exists between the nanomaterial and
the receptor molecules. Carbon nanotube field-effect transis-
However, there have been limited demonstrations utilizing
graphene for the detection of enzymes, with most demonstra-
tions relying on antibody-mediated recognition of their target
[17,31,32]
proteins.
Enzyme detection is an especially clinically
tors (CNT-FETs) currently lead the field and are one of the most
relevant issue as it has been reported that the enzyme levels
found in vivo are strongly correlated with the diagnosis, prog-
nosis, and treatment of many diseases such as HIV, cancer, and
diabetes. Therefore, to improve our ability to detect enzyme
levels and to enhance the clinical potential of graphene-based
biosensors, it is possible to take advantage of graphene’s sensi-
tivity to changes in electrical conductance in order to detect the
enzymatic cleavage of target peptides that are assembled on the
sensor surface.
extensively developed systems.[
24–29]
This is due to the excellent
electrical properties and the relatively well-developed surface
chemistry of CNTs. However, certain limitations do exist with
CNT-based electrical detection devices. These include inconsist-
encies in the assembly of CNTs resulting in device unreliability,
limits in its surface area, and a high-cost of fabrication.
An alternative nanomaterial, which has the potential to over-
come the aforementioned limitations of CNTs, is graphene.
Graphene consists of a single-atom-thick planar sheet of carbon
atoms that are arranged in a perfect honeycomb lattice and has
Herein, we report a novel approach for the detection of
enzymes using a graphene-nanoparticle (NP) hybrid biosensor.
This system can detect the activity of an enzyme and determine
its concentration by measuring the change in electrical hys-
teresis that results from the interaction of the target enzyme
with its corresponding substrate. In addition, we demonstrate
its ability to achieve the sensitive and selective detection of the
target enzyme. Similar electrical hysteretic behavior in carbon
nanomaterial-based devices with gate have been observed previ-
ously and is presumably caused by charge transfer from neigh-
boring adsorbates such as water molecules or from charge
Dr. S. Myung, Dr. C. Kim, A. Solanki, Prof. K.-B. Lee
Department of Chemistry and Chemical Biology
Institute for Advanced Materials
Devices, and Nanotechnology (IAMDN), Rutgers
The State University of New Jersey,
Piscataway, NJ 08854, USA
E-mail: kblee@rutgers.edu
http://rutchem.rutgers.edu/∼kbleeweb/
P. T. Yin, Prof. K.-B. Lee
[
18]
injection into the trap sites on the dielectric substrate. How-
ever, this is the first report of NP-mediated electrical hysteresis
in a graphene sensor. Moreover, the utilization of a hysteresis
change for sensing is advantageous because the immobilization
of specific enzymes on the graphene surface, which is required
for field effect transistor-type sensors, is unnecessary in our
system allowing us to preserve the superb electrical properties
of graphene. The use of hysteresis as a detection method also
allows graphene-based devices to overcome the low on-off ratio
caused by the zero band gap of graphene. In the majority of
Department of Biomedical Engineering
Rutgers, The State University of New Jersey
Piscataway, NJ 08854, USA
P. I. Reyes, Prof. Y. Lu
Department of Electrical and Computer Engineering
Rutgers, The State University of New Jersey
Piscataway, NJ 08854-8058, USA
J. Park, Prof. K. S. Kim
Center for Superfunctional Materials
Department of Chemistry
Pohang University of Science and Technology
Pohang, Korea
[
13–17]
previous studies,
a Dirac shift or a conductance change
at the fixed gate voltage that is associated with this Dirac shift
when the specific target material was attached selectively on the
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DOI: 10.1002/adma.201202961
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graphene channel was utilized. When using only a Dirac shift
for sensing, non-specific binding such as the physical attach-
ment of positively- or negatively-charged materials can also
result in a Dirac voltage shift or conductance change. However,
in our process, even though nonspecific binding on the channel
occurred, the amount of electrical hysteresis of our devices did
not change. Only when these devices were exposed to a specific
condition, did the hysteresis phenomenon start to decrease. Spe-
cifically, our graphene-NP hybrid biosensor is composed of gold
nanoparticles (AuNPs) that are assembled on a graphene sur-
face using a functional polypeptide linker, which concomitantly
functions as the substrate for the target enzyme. Proteolysis of
the functional polypeptide linker by the target enzyme leads to
disassembly of the NP layer and a measurable shift in electrical
hysteresis that can be visualized in the gate sweep curve. This
can then be used to quantify the enzyme concentration. AuNPs
were chosen because they are able to store charge, resulting in
hysteresis upon their release from the device surface.[ In our
case, as a proof-of-concept, we detected the activity of Carbox-
ypeptidase B in order to demonstrate the high achievable sen-
B, followed by a layer of AuNPs that have a negative charge on
their surface, and a peripheral ionic liquid, which functions
as the gate dielectric (Figure 1b(ii-iv)). Introduction of Carbox-
ypeptidase B would then initiate the hydrolysis of the specific
enzyme substrate, which in this case is the functional polypep-
tide linkers. The degradation of the polypeptide linkers result in
the release of AuNPs, which causes a hysteresis shift that can
be measured using a silver gate electrode. The range of hys-
teresis shift has been found to be proportional to the enzyme
concentration and can therefore be used to determine the level
of enzyme in the solution.
To fabricate the graphene-NP hybrid biosensor, the source
and drain Au electrodes were first fabricated on a SiO2 sub-
strate via the lift-off process. Graphene was deposited and pat-
[
2]
terned as previously described (See Supporting Information)
in order to form graphene channels between the Au electrodes
(Figure 1b(i)). Before assembling the AuNPs and the func-
tional polypeptide linker molecules on the graphene surface,
the surface was made hydrophilic through functionalization
with 4-(pyren-1-yl)butanal, which occurs via π−π interactions.
Specifically, functionalization was achieved by incubating the
device in a methanol solution (1:500) containing 4-(pyren-1-yl)-
butanal (See Supporting Information for synthesis) for 30 min-
utes (Figure 1b(ii)). Next, the aldehyde groups of 4-(pyren-1-yl)
butanal were coupled to the amine groups of the functional
polypeptide linker through reductive amination for 6 hours.
Once the graphene device, functionalized with polypeptide
linker, was placed in the AuNP solution, the negatively charged
AuNPs were able to self-assemble on the positively charged
functional polypeptide linker molecules via electrostatic inter-
actions (Figure 1b(iii)). Specifically, this is a result of the
remaining positively charged amine groups on the functional
polypeptide linkers. Finally, with the AuNPs
18]
[
30]
sitivity and selectivity of our graphene-NP hybrid biosensor.
This enzyme has been reported as a predictor of severe acute
pancreatitis, which is an acute inflammatory condition of the
pancreas. However, it is important to note that the reported bio-
sensing method can be extended to detect any number of key
enzymes, pH changes, and UV with the appropriate modifica-
tion to the functional linker substrate.
A schematic of our graphene-NP hybrid biosensor is depicted
in Figure 1a. The device consists of gold source and drain elec-
trodes that are connected by graphene channels (Figure 1b(i)).
The graphene channels are coated with a functional polypeptide
linker that is specific for our target enzyme, Carboxypeptidase
assembled on the graphene surface, we were
able to detect the electrical hysteresis and
gate effect of our device by utilizing an ionic
liquid as the gate dielectric.
The fabrication method developed in our
graphene-NP hybrid biosensor has several
advantages over conventional graphene-
based biosensors. First, we were able to fab-
ricate high performance graphene-NP hybrid
devices that have a large surface area avail-
able for biodetection and a high conductivity.
This was accomplished at a high yield and
uniformity by combining the properties of
graphene with the nanoscopic features of the
[
2]
AuNPs. Specifically, these properties allow
for the development of biosensors that have
high achievable current levels sufficient for
the measurement of minute conductance
changes regardless of the channel length
between two electrodes in the devices (Figure 1
and Figure 2a1). As a result, our device pro-
vides a larger number of sites that can be
conjugated with receptor molecules (e.g.
antibody or functional polypeptide linker) in
comparison with previous biosensors that
used reduced graphene oxide. In addition,
Figure 1. Schematic of the device. (a) Graphene-nanoparticle hybrid devices for enzyme
sensing. (b) Fabrication process of the hybrid biosensor. Fabrication of the graphene channel
between the Au electrodes (i). Functionalization of the graphene surface with hydrophilic mol-
ecules (ii). Assembly of the functional peptide linker molecules and AuNPs on this polypeptide
layer (iii). Sensing target biomolecules by measuring the change in electrical hysteresis
(iv). (c) Chemical structure of the functional polypeptide linker molecule.
2
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these reduced graphene oxide-based biosensors were limited
polypeptide linker. Furthermore, the graphene device with func-
tional polypeptide linker is characterized by a gate curve that
is similar to that seen previously with no observable electrical
by their low conductivity and limited achievable channel length
4]
(
e.g. only a few micrometers).[ Second, the functionalization
[18]
of the graphene surface with our functional polypeptide linker
followed by assembly of AuNPs was accomplished using π−π
interactions (graphene-pyrene aldehyde) and electrostatic inter-
actions (polypeptides-AuNPs). This method prevents the need
to covalently bond the linker to the graphene surface, there-
hysteresis. However, when negatively charged AuNPs were
assembled, a positive Dirac voltage shift occurred resulting in
a field effect due to the negative charge of the AuNPs. Inter-
estingly, a large electrical hysteresis was achieved due to the
memory effect during gate sweep. The difference in the Dirac
voltage between the forward and backward gate voltage sweeps
was defined as the hysteresis gap (ΔVDirac). In terms of the
underlying mechanism, graphene has both holes and electrons
as carriers, and the density of states in graphene near the Fermi
level is low. Thus, it exhibited ambipolar behaviors of some
gating effects depending on the gate bias voltages. Note that,
in the case without AuNPs (pristine graphene FET or graphene
FET with only linker molecules) (Figure 3a(i) and 3a(ii)), the
graphene FETs exhibit only a small amount of hysteresis indi-
cating that our devices do not have many charge trap defects.
On the other hand, in the case of devices with AuNPs
(Figure 3a(iii)), a large hysteresis is evident in the gate sweep
curve due to the charge stored in the AuNPs.
[
21]
fore preserving the superb electrical properties of graphene.
Finally, our fabrication method allows for the precise control
of the number of graphene layers and can be used to deposit
a large quantity of graphene on different substrates such as
SiO , metallic surfaces, or glass due to the flexible and stretch-
2
able properties of graphene. The devices can also be fabricated
with conventional microfabrication techniques, and the entire
process can be accomplished using only conventional microfab-
rication facilities. Therefore, our method allows us to fabricate
graphene-based devices with present Si-based microelectronics,
and would be readily accessible to the present semiconductor
industry. Collectively, our study presents a novel method that
avoids the use of graphene oxide for the development of biosen-
sors and provides a novel approach that combines the advan-
tages provided by graphene and inorganic nanoparticles (e.g.
metallic and semiconductive NPs) with the area of biosensing.
Confirmation of graphene deposition followed by AuNP
assembly in the junction between the Au electrodes was obtained
using scanning electron microscopy (SEM) (Figure 2A1). The
image shows that the two Au electrodes, separated by a 50 μm
gap, are connected by a graphene channel that is 420 μm wide.
In addition, after coating the graphene surface with the func-
tional polypeptide linker, we confirmed that the negatively
charged AuNPs (15 nm in diameter) successfully self-assembled
as a monolayer on the graphene surface at a high density
As stated above, Carboxypeptidase B has been proposed as
[
30]
a predictor for severe acute pancreatitis. This enzyme was
used in our proof-of-concept due to the fact that our functional
polypeptide linker contains a high number of lysines. There-
fore, it acts as a substrate for this enzyme, which specifically
cleaves basic amino acids such as lysine. When our graphene-
NP hybrid device was exposed to the enzyme solution, the func-
tional polypeptide linker degraded, resulting in the release of
AuNPs into the surrounding media. To characterize the effect
of the release of AuNPs from the graphene surface on the elec-
trical hysteresis of this device, we measured the ΔVDirac as a
function of time using a 1 μM enzyme solution (Figure 4b). We
found that increasing the exposure time of the enzyme resulted
in an expected decrease in NP density (Figure 2(c1-c6)). This
decrease in NP density resulted in a decrease in the measured
ΔVDirac, which became saturated after 50 hours. Furthermore,
due to the fact that the release of functional polypeptide linker
caused the detachment of NPs from the graphene surface,
the ΔVDirac was very similar to the release profiles found for
(Figure 2A4 and 2B2). This is mainly due to the large positive
charge and the two-dimensional planar structure of the func-
tional polypeptide linker. The negatively charged surface of the
AuNPs also ensures that additional NPs are not assembled onto
the underlying functional polypeptide linker. Therefore, the dep-
osition of a single layer of AuNPs with high density was achieved
on the graphene surface. It was also found that the complete dis-
assembly of AuNPs was achieved after the device was exposed
to a 1 μM Carboxypeptidase B enzyme solution for 50 hours
[23]
polypeptide films.
The sensitivity of our graphene-NP hybrid sensor with func-
tional polypeptide and NPs was determined by measuring the
ΔVDirac as the concentration of Carboxypeptidase B in phos-
phate buffered saline (PBS) solution was varied from 10 nM to
10 μM (Figure 3b). Representative sensing data shows that the
difference between the ΔVDirac for 10 nM and 100 nM solutions
of Carboxypeptidase B after 3 hours of exposure was statisti-
cally similar. However, a statistically significant difference in
ΔVDirac was observed when the concentration of Carboxypepti-
dase B was increased to 1 μM. Furthermore, as the concentra-
tion was subsequently increased, a concentration dependent
decrease in the ΔVDirac of our sensors was observed. Figure 4a
illustrates the sensitivity (ΔVDirac) of our biosensors as a func-
tion of the enzyme concentration. For these experiments, we
used ten graphene junctions for each step and repeated sensing
experiments ten times demonstrating the repeatability of our
results. From these results, we can conclude that the lowest
concentration that can be detected using the current functional
(
Figure 2C6) or a 10 μM solution for 3 hours (Figure 2D4).
The gating effect of our device was measured by employing
an ionic liquid as the gate dielectric and a silver gate electrode.
The use of an ionic liquid, which has a high ionic concentra-
tion, as the gate dielectric allowed the thickness of the diffusion
layer to be considered negligible. It was found that the effect
was similar to that observed in graphene transistors that have
ambipolar conduction and no electrical hysteresis (Figure 3a
upper graph).[ In our case, the top-gate bias was swept with
a ∼0.01 V/sec sweep speed under a source-drain bias of 0.1 V.
An ambipolar field effect was also observed with a Dirac point
voltage (VDirac) of 0.25 V, where charge carriers change polarity.
After functionalization of the graphene surface with our func-
tional polypeptide linker, it was observed that the gate curve
shifted to a lower voltage level. This negative Dirac voltage
shift is approximately 0.05 V and we hypothesize that this is
due to the positively charged amine groups of the functional
22]
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Figure 2. Characterization data of the biosensor using AFM and SEM. (A1) A SEM image of the junction between the Au electrodes, shows a graphene
channel (dark region) between two Au electrodes (bright area). AFM topography images of (A2) the graphene surface before the assembly of NPs and
(A3) the graphene surface after the functionalization with linker molecules. (A4) An AFM image of AuNPs on the graphene surface. The height profile
of the AFM image is shown below the image. (B1) An AFM topography image of the graphene surface before the assembly of NPs. (B2) An AFM image
of AuNPs on the graphene surface. The height profile of the AFM image is shown below the image. (B3) An AFM topography image of graphene after
exposure to a 10 μM enzyme solution for 6 hours. (C1)-(C6) AFM topography images of graphene after exposure to a 1 μM enzyme solution for 30 sec
to 50 hours, respectively. (D1)-(D4) AFM images after exposure to a 10 nM, 100 nM, 1 μM and 10 μM enzyme solution for 3 hours, respectively.
polypeptide linker is approximately 1 μM and it was found that
this concentration correlates with a decrease in the change
ratio of ΔVDirac (59.2%). The results also confirm that there is a
direct relationship between the ΔVDirac and the concentration of
enzyme. Specifically, similar to the non-linear behavior of other
graphene-based sensors, the responses of our sensor increase
non-linearly with an increase in the enzyme concentration
[
13,24–26]
from 10 nM to 10 μM.
Finally, we tested the selectivity of
this biosensor towards the target enzyme by using control sam-
ples such as PBS and different proteins such as bovine serum
albumin (BSA). Results showed that PBS and BSA did not react
with the functional polypeptide substrate (Figure 4b).
4
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Figure 3. Measurement of the hysteresis change of the device after exposure to enzyme solutions. (a) ΔVDirac change and gate effect of an initial
graphene transistor (i), after the functionalization of graphene surface with functional polypeptide linker molecule (ii), after NP assembly on the
graphene surface (iii), and after exposure to 1 μM enzyme solution for 30 sec, 5 min, 1 hr and 50 hrs (iv ∼ vii). (b) Changes in ΔVDirac of graphene
device (i), after functionalization with linker molecules (ii), after NP assembly (iii), after the exposure to Carboxypeptidase B in PBS solution with
the concentration of 10 nM (iv), 100 nM (v), 1 μM (vi) and 10 μM (vii) for 3 hours.
Our findings show that quantifying the change in electrical
hysteresis of a graphene-NP hybrid biosensor is a viable method
to measure enzyme activity sensitively and selectively. However,
it is apparent that the critical factor that determines this sensi-
tivity and selectivity is the functional polypeptide linker whose
degradation results in the release of AuNPs and subsequent
shift in hysteresis. Therefore, the performance of our sensor, in
terms of its sensitivity and selectivity, can be further improved
by modifying the functional polypeptide linker and changing
the properties of the metallic nanoparticle system. In the cur-
rent study, our novel graphene-based biosensor with functional
polypeptide linker and NPs can confer a sensitivity of up to
ability to sense concentrations of Carboxypeptidase B down
to the μM-scale, which is comparable to previously reported
[
15–17]
graphene-based biosensors.
More importantly, to the
knowledge of the authors, this is the first demonstration of a
graphene-based biosensor that utilizes a hysteresis change
resulting from metallic NPs assembled on a graphene surface
for enzyme detection. The current configuration developed in
our graphene-NP hybrid biosensor holds a number of advan-
tages over other graphene-based devices established to detect
enzymes and other biomolecules. First, AuNP-mediated elec-
trical hysteresis is a phenomenon that has not been studied in
graphene-based electrical detection methods, even though it
can overcome the limitations faced by current graphene-based
sensing approaches. Second, the novel method of detection
reported here provides an alternative to the low on-off ratio
caused by the zero band gap of graphene. Third, in terms of
the components used in our device, the functional polypeptide
linker acts as the key-sensing component and can be modified
to enhance sensitivity and selectivity for any target enzyme or
to confer pH and UV sensitivity. In addition, by combining the
properties of graphene with inorganic NPs, our device provides
1
μM. However, to improve sensitivity and selectivity, we are
synthesizing and screening a variety of functional polypep-
tide linker molecules that will be self-assembled with different
metallic NPs of varying compositions and sizes.
In summary, we have successfully developed a novel
graphene-NP hybrid biosensor that uses an electrical hyster-
esis change to detect the enzymatic activity and concentra-
tion of a target enzyme (e.g., Carboxypeptidase B). Our results
indicate that our novel graphene-NP hybrid biosensor has the
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ppm downfield to TMS (δ = 0 ppm) were observed:
1
H NMR (600 MHz, DMSO-D ): δ 0.91(t, 3H), δ
6
1
(
.27(m, 2H), δ 1.77(m, 2H), δ 3.85(s, 3H), δ 4.16
t, 2H), δ 7.67(s, 1H), δ 7.74(s, 1H), δ 9.08(s, 1H).
Synthesis of 4-(pyren-1-yl)butanal: 10 mL of
methylene chloride was added to 1.097 g (4 mmol)
of 4-(pyren-1-yl) butyl alcohol and stirred vigorously
at room temperature for 10 min. The reaction was
followed by the addition of 1.292 g (1.5 Equilibrium)
of PCC (Pyridinium chlorochromate) in 10 mL
methylene chloride and the reaction mixture was
stirred for 2 h. The reaction mixture was then
diluted with 5 volumes of anhydrous ether (100 mL)
and washed with 1:1 brine: water, saturated aq.
Figure 4. Enzyme detection using the graphene-nanoparticle hybrid sensor (a) Sensitivity Na SO solution, and brine, respectively, dried over
2
3
(
ΔVDirac) of the graphene sensor as a function of enzyme concentration [n = 10]. (b) ΔVDirac
change under various periods of exposure to a 1 μM solution of Carboxypeptidase B in PBS aldehyde. The crude product was purified by silica
black square), a 1 mM PBS solution (blue triangle) and 1 mM solution of BSA in the PBS (red gel flash chromatography (hexanes: ethyl acetate,
anhydrous Na SO4 and concentrated to give the
2
(
circles) [n = 10].
9:1, rf: 0.4) to yield 1.06 g (97%). Most of the crude
products were very clean and could be used directly
for further applications.
a larger number of sites that can be conjugated with receptor
molecules and a higher achievable current level. Fourth, the
methods used to fabricate the device allow for the homogenous
deposition of graphene, resulting in uniform electrical prop-
erties. Finally, the use of conventional microfabrication tech-
niques allows for compatibility with facilities found in device
industries and allows mass production of our device. Therefore,
the results and developed methods presented here warrant the
further study of graphene-NP hybrid biosensors for the sensi-
tive and selective detection of enzymes and other biomolecules.
Supporting Information
Supporting Information is available from the Wiley Online Library or
from the author.
Acknowledgements
This work was supported by the NIH Director’s Innovator Award
[
(1DP20D006462–01), K.-B. L.] and the N. J. Commission on Spinal Cord
Injury grant [(09–3085-SCR-E-0), K.-B. L.]. We are grateful to Dr. Barman
for his valuable suggestions and scientific comments with regard to the
manuscript.
Experimental Section
Metal deposition and measurement of graphene devices: For the
electrode fabrication, the photoresist was first patterned on the
substrate. Ti/Au (10/30 nm) was then deposited on the substrate and
the remaining photoresist was removed with acetone for the lift-off
process. A Keithley-4200 semiconductor parameter analyzer was used
for measurement and data collection.
Received: July 24, 2012
Published online:
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water until no chloride (Cl ) anion in the water layer could be detected
[
[
by AgNO3 solution. Dichloromethane was evaporated and 1-butyl-3-
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6
colorless or slightly yellow viscous liquid after drying in a vacuum oven.
[
1
-chlorobutane, 1-methylimidazole, ethyl acetate and dichloromethane
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