Nano Res
1
Oxide-on-graphene field effect bio-ready sensors
Bei Wang1†
, Kristi L. Liddell2†
, Junjie Wang1, Brandon Koger
1, Christine D. Keating
2, and Jun Zhu
1,3 ()
Nano Res., Just Accepted Manuscript • DOI: 10.1007/s12274-014-0489-9
http://www.thenanoresearch.com on May 3, 2014
© Tsinghua University Press 2014
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Nano Research
DOI 10.1007/s12274-014-0489-9
1
TABLE OF CONTENTS (TOC)
Oxide-on-graphene field effect bio-ready sensors
Bei Wang, Kristi L. Liddell, Junjie Wang, Brandon
Koger, Christine D. Keating, and Jun Zhu*
The Pennsylvania State University, USA
This work reports on the fabrication, functionalization and operation of
a novel oxide-on-graphene field effect bio-ready sensor. pH sensing is
demonstrated.
Provide the authors’ website if possible.
Jun Zhu, http://www.personal.psu.edu/jxz26/zhulab/
2
Oxide-on-graphene field effect bio-ready sensors
Bei Wang1†, Kristi L. Liddell2†, Junjie Wang1, Brandon Koger1, Christine D. Keating2, and Jun Zhu1,3 (*) 1 Department of Physics, The Pennsylvania State University, University Park, Pennsylvania 16802, USA 2 Department of Chemistry, The Pennsylvania State University, University Park, Pennsylvania 16802, USA 3 Materials Research Institute, The Pennsylvania State University, University Park, Pennsylvania 16802, USA † These authors contributed equally to this work. Received: day month year / Revised: day month year / Accepted: day month year (automatically inserted by the publisher) © Tsinghua University Press and Springer-Verlag Berlin Heidelberg 2011
ABSTRACT
Electrical detection schemes using nanoscale devices offer fast and label-‐‑free alternatives to biosensing
techniques based on chemical and optical interactions. Here we report on the design, fabrication, and
operation of oxide-‐‑on-‐‑graphene ion-‐‑sensitive field effect sensor arrays using large-‐‑area graphene sheets
synthesized by chemical vapor deposition. In this scheme, HfO2 and SiO2 thin films are deposited atop the
graphene sheet and play the dual role of the sensing interface, as well as the passivation layer protecting the
channel and electrodes underneath from direct contact with the electrolyte. We further demonstrate the
functionalization of the SiO2 surface with 3-‐‑aminopropyltrimethoxysilane (APTMS). The oxide-‐‑on-‐‑graphene
sensors operate in solution with high stability and a high average mobility of 5000 cm2/Vs. As a proof of
principle, we demonstrate pH sensing using the bare or the APTMS-‐‑functionalized SiO2 as the sensing
surface. The measured sensitivities, 46 mV/pH and 43 mV/pH respectively, agree well with existing studies.
We further show that by applying the solution gate voltage in pulse, the hysteresis in the transfer curve of
the graphene transducer can be eliminated, greatly improving the ionic potential resolution of the sensor.
Nano Res DOI (automatically inserted by the publisher) Research Article
Jun � 2/16/14 4:07 PM
Deleted: bio
Jun � 2/16/14 4:13 PM
Deleted: The latter enables further attachment ofto protein and nucleic acid
molecules
Jun � 2/16/14 4:13 PM
Deleted: . This design utilizes both the excellent transport characteristics of the
graphene transducer and the
well-‐‑established linker chemistry of the
SiO2 surface to achieve specific binding
and detection of biomolecules.
Jun � 2/16/14 4:12 PM
Deleted: 3-‐‑aminopropyltrimethoxysilane
(
3
These experiments demonstrate the potential of oxide-‐‑on-‐‑graphene ion-‐‑sensitive field effect sensors in
on-‐‑chip, label-‐‑free and real-‐‑time biosensing applications.
KEYWORDS graphene, biosensor, pH sensor, ISFET, APTMS, label-‐‑free
1. Introduction
Sensors are essential to biomedical diagnosis. Compared to traditional methods such as fluorescence,
surface plasmon resonance, and bioassays, sensors based on electrical detection schemes can potentially be
faster, more cost effective, and require less specialized equipment. Ion-‐‑sensitive field effect transistors
(ISFETs), for example, work by converting charge accumulation caused by the binding of biomolecules on a
sensing surface into a potential signal, which is then detected by the conductance change of the FET
transducer [1]. Such sensors can be readily integrated into a multiplexing system to deliver low-‐‑cost, on-‐‑chip
biomedical diagnosis [2-‐‑4], or used to study the hybridization and binding kinetics of DNA and proteins
[5-‐‑10].
Until recently, ISFET technology primarily focused on silicon transistors [1, 7, 8, 11]. The sensitivity of
bulk devices is ultimately limited by its bulk carrier mobility of ~500 cm2/Vs. More recent efforts explore field
effect transistors (FETs) based on low-‐‑dimensional nanostructures, such as nanowires and carbon nanotubes,
because of their size, large surface-‐‑to-‐‑volume ratio, and potentially higher sensitivity. Although high
performances have been demonstrated [12-‐‑15], much work needs to be done to address practical issues such
as uniformity, stability and scalability before applications can be developed [16, 17]. Graphene FET may be a
———————————— Address correspondence to Jun Zhu, [email protected]
Jun � 2/16/14 4:14 PM
Deleted: romise
4
good candidate for biosensing because of its excellent carrier mobility and the availability of low-‐‑cost
large-‐‑scale synthetic methods [18-‐‑21]. Indeed, several recent studies report using graphene FETs as a pH
sensor [22-‐‑25] and for the detection of protein and DNA [23, 26, 27]. However, the underlying sensing
mechanism of a graphene FET is not clear. The reported pH sensitivities of as-‐‑grown graphene sheets vary
from 6—99 mV/pH [22-‐‑25, 28]. Fu et al. demonstrates that a pristine graphene sheet is insensitive to H+
concentration change in solution [28] and points to the role of uncontrolled extrinsic imperfections, such as
defects and contaminations, in the sensing process. These complications require further studies to clarify.
Furthermore, a pristine graphene surface precludes specific binding and detection. Functionalization,
however, severely reduces the carrier mobility in graphene. For example, the introduction of 1013/cm2 sp3
centers (e. g. C-‐‑NH2 to bind to protein and DNA) will reduce the mobility to less than 100 cm2/Vs,
compromising the advantage of graphene FETs [29-‐‑31]. In comparison, SiO2 has been widely used as the
dielectric layer and sensing surface in ISFETs made from bulk silicon to nanowires because of the
well-‐‑established silanization chemistry that enables the immobilization of specific bioprobes and targets [32].
The ion sensitivity of a SiO2 surface is well understood by the site dissociation model [33, 34].
In this work, we demonstrate the fabrication and operation of a novel oxide-‐‑on-‐‑graphene bio-‐‑ready
sensor. In this scheme, the graphene channel is passivated by a SiO2/HfO2 (25 nm/20 nm) double oxide layer,
which uses the top SiO2 layer as the sensing/immobilization surface and the high-‐‑quality HfO2 layer to
protect the channel from the solution. This design preserves the high carrier mobility of the graphene
transducer, which averages 5000 cm2/Vs in our devices. We demonstrate the functionalization of the SiO2
surface with 3-‐‑aminopropyltrimethoxysilane (APTMS), as a pathway to the immobilization of proteins and
DNA [35]. The devices operate stably and reproducibly in solution. Bare and APTMS-‐‑functionalized devices
5
are sensitive to the pH value of phosphate buffer saline (PBS) solutions ranging from 4-‐‑9. The observed
sensitivity of 46 mV/pH and 34 mV/pH respectively are in good agreement with results obtained on similar
surfaces [1, 33, 34, 36]. Operating the solution gate in pulse mode eliminates the hysteresis in the transfer
curve of the graphene FET, which is a common issue in nanostructure-‐‑based biosensors. Our experiments
show the potential of this new graphene-‐‑based device design in on-‐‑chip sensing applications.
2. Experimental
The pH sensing study uses 0.010 M PBS (0.010 M potassium chloride and 0.001 M sodium phosphate, pH
4.1, 6.0, 6.9, or 8.8); chemicals are purchased from Sigma Aldrich. HPLC grade water for buffer preparation
is purchased from EMD Chemicals.
APTMS functionalization is carried out in 3% 3-‐‑aminopropyltrimethoxysilane (APTMS, TCI America) in
200 proof ethanol (Koptec) for 1 hour. Excess APTMS is removed with an ethanol rinse followed by a 10
minute cure at 110oC.
pH sensing measurements are performed by covering the device channel area in a droplet of PBS
solution of the desired pH, avoiding contact between the liquid and the electrodes. A silicone ring is
sometimes used to confine the droplet. Transfer curves are obtained at each pH value in the course of ~20
minutes. After each pH measurement, the chip is thoroughly rinsed in DI water followed by the PBS solution
of the next desired pH value before the next set of measurements is begun.
3. Results and discussion
The graphene sheets used in this study are synthesized by a low-‐‑pressure chemical vapor deposition
Kristi Liddell� 3/25/14 2:19 PM
Deleted: romise
6
(CVD) technique on copper foil [18]. Figure 1(a) shows a scanning electron microscope (SEM) image of
as-‐‑grown graphene, showing a full monolayer coverage with ~5% multi-‐‑layer islands. The sheets are
transferred to 290 nm SiO2/highly doped Si substrates using a polymer-‐‑assisted wet transfer method [18]. A
typical Raman spectrum of transferred graphene is given in Fig. 1(b), where a small ID/IG ratio of less than 0.1
indicates high-‐‑quality growth [18]. The transferred sheet is annealed in Ar/H2 at 450°C for 2 hours. After
annealing, the graphene surface appears flat, continuous and mostly free of polymer residue, with small
amount remaining at the wrinkles and folds of the sheet.
Two-‐‑terminal graphene FETs of dimensions 2 µμm (width) × 4 µμm (length) are fabricated using optical
lithography, reactive ion etching and metal deposition. We then deposit 20 nm of HfO2 on the whole wafer
using atomic layer deposition and recipes previously established by our group [37], followed by the electron
beam evaporation of 25 nm of SiO2. An optical micrograph of a sensor array and an SEM image of a channel
are shown in Fig. 2.
Figure 3(a) shows a schematic of the oxide-‐‑on-‐‑graphene sensor operating in solution, using both the
doped silicon back gate and the solution top gate. A small droplet of deionized water or PBS solution forms
the solution gate above the channel area (see Fig. 2). Care is taken to ensure that the droplet does not cover
the contact pads of the device being measured. A tungsten electrode and an Ag/AgCl reference electrode are
inserted into the droplet to apply Vapp and read Vsg, the potential of the solution, respectively. The use of the
reference electrode is necessary to eliminate spurious changes in the characteristics of the sensor due to the
variation of the potential drop at the tungsten/solution interface. Because of the conformal growth of the
HfO2 film, the channel is well protected from the solution and the solution gate operates with a small leakage
current of less than 100 pA.
7
Figure 3(b) plots the conductance of a graphene channel as a function of the back gate voltage Vbg (top
axis) and the solution gate voltage Vsg (bottom axis) using different symbols. The two axes are scaled and
offset with respect to one another so that the two measured traces overlap. The device is ambipolar with a
small amount of unintentional doping. The back gate sweeps are typically hysteretic with a separation of
ΔVbg ~7 V between the two charge neutrality points. Only one sweep direction is shown here. The solution
gate hysteresis ΔVsg is generally less than 100 mV. The scaling ratio γ between the two gate voltages Vsg and
Vbg in Fig. 3(b) gives the ratio between the capacitance of the solution gate, Csg and that of the back gate Cbg,
i.e., Csg = γ Cbg. Measurements on 11 devices yield γ = 11.3 ± 1.1, from which we determine Csg= (127 ± 12)
nF/cm2 using a back gate capacitance Cbg= 11.2 nF/cm2. This result can be understood by noting that Csg
consists of the capacitance of the solution CDL, and the capacitance of the double oxide layer (20 nm HfO2 and
25 nm SiO2) Cox in series:
.
In our experiment, CDL, which scales with the ionic concentration of the solution M as CDL~M1/2[5], is
much larger than Cox such that Csg≈ Cox. We determine Cox to be ~120 nF/cm2 through direct capacitance
measurement, which agrees very well with Csg obtained here.
We determine the field-‐‑effect mobility µFE of the graphene FETs by fitting the G(Vbg) curve to the
charged-‐‑impurity scattering model, following Eq. (2) of Hong et al [38]. The fit describes data very well, as
shown in Fig. 3(b). The field effect mobility µFE averages approximately 5000 cm2/Vs for both electrons and
holes, indicating the high quality of our devices among CVD-‐‑grown graphene [21]. The high µFE and Csg
together lead to a high normalized transconductance of gm= µFE·∙ Csg = 0.6 mS/V. In devices measured here, gm
€
Csg =CDLCox
CDL +Cox
8
is reduced by a factor of a few due to the contact resistance, which is around 5 kΩ in our devices. The contact
resistance can potentially be reduced by a few orders of magnitude by using gentle oxygen plasma etching
and annealing [39] to further increase gm.
In field effect devices made from nanomaterials such as nanotubes, nanowires, and graphene, hysteretic
transfer curves caused by interfacial changes and adsorbates are a common problem [13, 23, 40]. Adopting
pulsed gate sweep techniques (Fig. 4(b)) applied to carbon nanotube FETs [41, 42], we can suppress the
hysteresis in our graphene FETs completely. Examples of hysteretic and hysteresis-‐‑free G(Vsg) traces of the
same device are shown in Fig. 4(a). The elimination of hysteresis eliminates a source of uncertainty in the
operation of nanostructured sensors and greatly improves their long-‐‑term stability.
Next we demonstrate the proof-‐‑of-‐‑principle operation of the oxide-‐‑on-‐‑graphene sensor by measuring its
response to the pH of PBS solutions. Both bare SiO2 and APTMS-‐‑functionalized devices are tested and their
sensing performance evaluated in Figs. 5 and 6. The test protocol is described in the experimental. Briefly, we
measure the G(Vsg) curve of the graphene FET while it is immersed in a PBS solution droplet of a fixed pH
value. The device is thoroughly rinsed with de-‐‑ionized water prior to exposure to another droplet of a
different pH. Measurements of the same pH value are repeated at different times to check the stability of the
sensor. The same test is done on many graphene FETs.
Figure 5 shows the pH response of a bare FET. As the pH value of the PBS solution increases, the G(Vsg)
curve retains its shape but shifts toward positive Vsg. The inset of Fig. 5 plots the position of the Dirac point
VD vs pH. The error bars represent the spread of VD from repeated measurements in the course of 6 hours.
Measurements are reproducible after weeks, demonstrating excellent stability of the sensors. In the range of
pH = 4.1 to 8.8, VD (pH) is well described by a linear fit, with a slope of 43 mV/pH. Averaging many devices,
we obtain a voltage sensitivity of (46 ± 8) mV/pH. These results are in good agreement with literature results
9
and can be well understood by the site-‐‑dissociation model developed for oxide surfaces [1, 33, 34]. Briefly,
the pH of the solution affects the protonation and deprotonation of the oxide surface, leading to a change in
the potential drop ψ0 at the solution/oxide interface, which causes the G(Vsg) curve to shift in Vsg. In the case
of SiO2, increasing pH increases the presence of negative charges on the SiO2 surface, which results in a
larger ψ0, and a shift in G(Vsg) towards positive Vsg. This is exactly what we observed. In the linear regime of
G(Vsg) in Fig. 5, the conductance changes by 4.2 µS/pH. This is a large conductance change among
nanostructured sensors [4, 12] and can be further increased by reducing the oxide thickness (presently 45 nm)
and increasing carrier mobility (presently 5000 cm2/Vs).
In Table I, we compare the voltage sensitivity of our oxide-‐‑on-‐‑graphene sensor with other graphene FET
pH sensors reported in the literature. Existing studies show a large variation ranging from 6 to 99 mV/pH using
as-‐‑grown graphene obtained from different methods. This large variation makes graphene pH sensors ill suited
for applications. The underlying sensing mechanisms are not well understood at the present, but the presence
of defects, contaminations and unintentional chemical functionalizations has been suggested to play key
role[22, 28]. In contrast, the sensitivity of our oxide-‐‑on-‐‑graphene sensor is given by the well-‐‑characterized oxide
surface and is well understood. Indeed, ISFET based pH sensors are commercially available and used in
applications requiring continuous readings such as environmental monitoring[43]. They are also a crucial
component of a CMOS integrated non-‐‑optical genome sequence scheme demonstrated recently[8]. The
oxide-‐‑on-‐‑grahpene design can potentially take advantage of the CMOS architecture already developed for
silicon ISFETs while using higher mobility graphene as the FET transducer.
We next show that oxide-‐‑on-‐‑graphene FETs can be readily functionalized with amine groups, which are
commonly used to immobilize and recognize biomolecular probes and targets [7]. The top SiO2 surface of the
graphene FETs are functionalized with APTMS following procedures described in the experimental section.
Figure 6(a) shows the high-‐‑resolution X-‐‑ray photoelectron spectroscopy (XPS) spectra of the N-‐‑1s state on
bare and APTMS-‐‑functionalized SiO2 surfaces, respectively. The spectrum on functionalized surface exhibits
two prominent peaks at 399.9 and 401.9 eV whereas no such signal is detected on the bare SiO2 surface. These
peaks correspond to NH2 and NH3+ groups, respectively, with the latter due to the presence of water [44].
Atomic force microscopy shows that most areas of the functionalized surface are covered by a thin film of
average thickness 0.7 nm (Fig. 6(b)), which is consistent with the height of one APTMS molecule [36]. In
10
some areas, partial coverage as well as local aggregation of APTMS molecules is also observed. Similar to the
bare oxide-‐‑on-‐‑graphene devices, G(Vsg) of APTMS-‐‑functionalized devices shift towards positive Vsg with
increasing pH, an example of which is shown in Fig. 6(c). The inset of Fig. 6(c) plots the pH-‐‑dependent Dirac
point VD of this device, from which we extract a slope of 28 mV/pH. Averaging many devices, we find the
pH sensitivity of APTMS-‐‑functionalized graphene FETs to be (34 ± 8) mV/pH. These results agree well with
previous studies [36]. The reduced pH sensitivity is expected, as the amine group is less amphoteric than the
Si-‐‑OH group and the coverage of the APTMS layer is also probably less dense than that of the SiO2. The
functionalization of the oxide-‐‑on-‐‑graphene sensor with APTMS paves the path to the immobilization and
detection of specific binding events of molecules, such as the hybridization of DNAs.
4. Conclusions
In summary, we have designed, fabricated and demonstrated the operation of a novel high-‐‑quality,
oxide-‐‑on-‐‑graphene field effect bio-‐‑ready sensor in solution. The use of thin oxide film as the sensing layer
preserves the high mobility of the graphene transducer and enables sensing specificity. The
oxide-‐‑on-‐‑graphene FETs function stably and reproducibly. As a proof of principle, we show that bare and
APTMS-‐‑functionalized SiO2 surfaces response to the pH of PBS solutions, with the sensitivity of 46 mV/pH
and 34 mV/pH respectively. Our studies open the door to using graphene-‐‑based electrical devices to
selectively detect the presence and binding events of molecules of interest.
Acknowledgements
We thank Xiahua Zhong and Wenchong Hu for helpful discussions on experimental setup. The synthesis
of graphene, device preparation and electrical measurements are supported by NSF NIRT grant No.
ECS-‐‑0609243, MRSEC grant No. DMR-‐‑0820404 and CAREER grant No. DMR-‐‑0748604. Wet chemistry and
11
chemical functionalization work was supported by the MSD Focus Center, one of six research centers funded
under the Focus Center Research Program (FCRP), a Semiconductor Research Corporation entity. The authors
acknowledge use of facilities at the PSU site of NSF NNIN.
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14
Figures
Figure 1. (a) SEM image of a continuous graphene film grown on copper substrate. Small, dark islands are multi-layers. (b) A typical
Raman spectrum of graphene transferred to SiO2/doped Si wafer showing a small ID/IG ratio of less than 0.1.
Figure 2. Optical image of one oxide-on-graphene field effect transistor array consisting of sixteen 2 µm × 4 µm two-terminal
graphene channels. The inset shows an SEM image of one such channel.
1µm
a
1500 2000 2500 1500 2000 2500
Raman Shift (cm-1)
Inte
nsity
(Arb
. Uni
t) b
1 mm
4µm
15
Figure 3. (a) Schematic drawing of the oxide-on-graphene FET operating in solution. The solution gate voltage is applied
through a tungsten electrode and read by an Ag/AgCl reference electrode. (b) The conductance of a graphene channel vs the
solution gate voltage Vsg (magenta solid triangle, bottom axis) and the backgate voltage Vbg (blue hollow square, top axis). The
two traces overlap well after scaling the Vbg and Vsg axes. The black solid line is a fit to the charged impurity model,
1/G=Rc+2/(neuFE+σres) following Eq. (2) of Ref. [38]. Fits to electrons and holes separately yield µFEe = 3900 and µFE
h = 5300
cm2/Vs respectively and Rce = 4700 Ω and Rc
h = 6300 Ω. σres= 0.17 mS.
Figure 4. (a) The conductance of a graphene channel vs the solution gate voltage Vsg with Vsg changed continuously (solid
magenta traces) and in pulse (blue hollow circles). The arrows indicate the sweeping directions of the magenta traces. The
time-varying pattern of the applied pulse is shown in (b) thigh = 25 ms, tlow = 75 ms
resubmission*
Vapp Vsg
- - - - - - - - + + + + + + + + +
Vbg
V
~Vds
a
60.0
80.0
100.0
120.0
-12 -8 -4 0 4
-1.5 -1.0 -0.5 0.0 0.5
Vbg(V)
Vsg(V)
120
100
80
60
-12 -8 -4 0
-0.5 0 0.5
b
G(µ
S)
W Ag/AgCl
290nm SiO2
25nm SiO2 20nm HfO2
p++ Si
S D
Vsg(V)
-0.6 -0.4 -0.2 0 0.2
G(µ
S)
thigh!
tlow!
a b
0.4 120
140
160
180
200
16
Figure 5. G(Vsg) curves of a graphene sensor with bare SiO2 surface in PBS solutions of different pH values as indicated in the
plot. From left to right: pH = 4.1 (blue), 6.0 (magenta), 6.9 (black), 8.8 (red). Inset: The Dirac point voltage VD as a function
of the pH value. The solid line is a linear fit with the slope of 43 mV/pH and a correlation coefficient greater than 0.99.
-0.6 -0.5 -0.4 -0.3 -0.2 -0.1
80
90
100
110
120
130 pH4 1st pH5.25 1st pH6 1st pH7 1st pH8.75 1st pH10
G(µ
S)
Vref(V)
4.1 6.0 6.9 8.8
Vsg (V)
G(µ
S)
17
Figure 6. (a) High-resolution XPS spectra of the N-1s state on bare (blue trace) and APTMS-functionalized SiO2 surfaces
(magenta trace). (b) AFM micrograph showing a differential height of 0.7 nm between areas covered by APTMS and where
the molecules were removed by contact mode scanning prior to acquiring this image. (c) G(Vsg) traces of an APTMS
functionalized sensor in PBS solution of different pH values. From left to right: pH = 4.1 (blue), 5.0 (magenta), 6.9 (black),
8.9 (red). Inset: The Dirac point voltage VD as a function of the pH value. The solid line is a linear fit with the slope of 28
mV/pH and a correlation coefficient of 0.97.
405 400 395
CP
S (A
rb. u
nit)
Binding energy (eV)
10
8
6
4
2
0
µm
1086420µm
-1.5
-1.0
-0.5
0.0
0.5
1.0
1.5
nm
a b NH3
+ NH2
0 2 4 6 8 10-0.50.00.51.01.52.02.53.03.5
Heigh
t (nm
)
X (µm)
-0.3 -0.2 -0.1 0.0 0.1 0.2 0.3
40.0
42.0
44.0
46.0
pH4 pH5 pH6 pH7 pH9
G (µ
S)
Vref (V)
4.1 5.0 6.9 8.9
z*
c
Vsg (V)
G(µ
S)
2 µm
18
Table I. Comparison of graphene pH sensors
Ang et al[22]
This work
Ohno et al[23]
Giacchetti et al[25]
Cheng et al[24]
Fu et al[28]
sensitivity (mV/pH)
99 46 27 24 18 6
graphene source epitaxial CVD exfoliation CVD exfoliation CVD
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