Ultrasensitive detection of Ebola matrix protein in a memristor mode
Bergoi Ibarlucea1,2 (), Teuku Fawzul Akbar1, Kihyun Kim3, Taiuk Rim3, Chang-Ki Baek3, Alon Ascoli4,
Ronald Tetzlaff4, Larysa Baraban1,2 (), and Gianaurelio Cuniberti1,2
1 Institute of Materials Science, Max Bergmann Center for Biomaterials, Technische Universität Dresden, Budapester Str. 27, Dresden
01069, Germany 2 Center for Advancing Electronics Dresden (CFAED), Technische Universität Dresden, Dresden 01069, Germany 3 Department of Creative IT Engineering, Pohang University of Science and Technology, Pohang 37673, Republic of Korea 4 Chair of Fundamentals of Electrical Engineering, Technische Universität Dresden, Mommsenstraße 12, Dresden 01069, Germany
Received: 16 April 2017
Revised: 2 June 2017
Accepted: 11 June 2017
© Tsinghua University Press
and Springer-Verlag Berlin
Heidelberg 2017
KEYWORDS
memristor biosensor,
capacitance,
honeycomb nanowires,
silicon nanowire field
effect transistor,
VP40 matrix protein,
Ebola detection
ABSTRACT
We demonstrate the direct biosensing of the Ebola VP40 matrix protein, using a
memristor mode of a liquid-integrated nanodevice, based on a large array of
honeycomb-shaped silicon nanowires. To shed more light on the principle of
biodetection using memristors, we engineered the opening of the current-minima
voltage gap VGAP by involving the third gap-control electrode (gate voltage, VG)
into the system. The primary role of VG is to mimic the presence of the charged
species of the desired sign at the active area of the sensor. We further showed
the advantages of biodetection with an initially opened controlled gap (VGAP ≠ 0),
which allows the detection of the lowest concentrations of the biomolecules
carrying arbitrary positive or negative charges; this feature was not present in
previous configurations. We compared the bio-memristor performance, in terms
of its detection range and sensitivity, to that of the already-known field-effect
transistor (FET) mode by operating the same device. To our knowledge, this is
the first demonstration of Ebola matrix protein detection using a nanoscaled
electrical sensor.
1 Introduction
Nanosensors are currently attracting attention as pro-
mising tools for biotechnology, e.g., as a miniaturized
diagnostics laboratory; being able to sense pH [1–4]
close to the Nernst limit; and to detect biomolecules
[5–7], viruses [8], and cell activities [9] with high
sensitivity. Label-free, rapid, and ultrasensitive probing
of biologically relevant analytes is possible at a
nanoscale and at low cost thanks to devices that rely
on current or voltage changes that are altered by
charge redistribution on the active surface or in the
Nano Research
DOI 10.1007/s12274-017-1720-2
Address correspondence to Bergoi Ibarlucea, [email protected]; Larysa Baraban, [email protected]
| www.editorialmanager.com/nare/default.asp
2 Nano Res.
surrounding environment [10]. Among the multiple
classes of electronic sensors, devices measuring the
electrical impedance [11–13] and field-effect transistors
(FETs) [14–18] have been shown to provide the best
detection limits without sacrificing the possibility to be
miniaturized and implemented as wearable, flexible
devices [19]. FETs are the most-studied configuration,
since their introduction by Piet Bergveld [20] in 1970
and technological and conceptual upgrade using silicon
nanowires by Charles Lieber [4] approximately a decade
ago. Several works have shown the large number of
applications and possibilities that FETs offer [21–25].
However, they reveal an important drawback; the
sensitivity is lost when samples with a high ionic
concentration are measured, owing to shortening of the
Debye length—the distance within which the devices
are sensitive [26].
During previous years, strong efforts have been
dedicated to overcome the limitation of FETs in higher
ionic strength solutions or to gain enough sensitivity
to allow detection by simply diluting the sample.
Multiple strategies have involved surface chemical
modifications such as the incorporation of biomolecule-
permeable polymer layers [27] or gold nanoparticles
with antibody fragments [28]. Others, for example, rely
on new analytical procedures, such as the analysis of
antigen-dissociation kinetics [29], the analysis of high
frequency impedance signals [30, 31], and the operation
of the device by dual gates [32].
An alternative strategy is to explore the distinct
nanoelectronic configurations of the sensing elements
and measurement modes. In this context, after the
theoretical prediction of the existence of the memristor
in 1971 by Leon Chua [33] and the association of his
ideas with a fabricated device in 2008 by Strukov et al.
[34], studies of this new type of circuit element have
been conducted [35, 36], including Sandro Carrara’s
demonstration of their application as a new type of
biosensor [37]. Tzouvadaki et al. [38] demonstrated
that a non-zero concentration of charged biomarkers
opens up the semi-logarithmic output curves (drain
current–source-to-drain voltage I–VSD loci), under
alternating current (AC) excitation. In other words, the
perturbations caused by the biomarker presence on
the sensing surface (Fig. 1(a)) lead to the violation of
one of the memristor signatures, namely, the exhibition
of pinched (i.e., zero gap) current–voltage loci [39].
The appearance of a gap between the minimum current
peak values and the voltage width that separates
them (voltage minimum gap, VGAP, Fig. 1(b)) has been
associated with larger energy requirements for the
transport of charge carriers compared to that required
for the bare genuine memristor with no gap. The gap
depends on the amount and the charge sign of the
biomarkers. This leads to the high sensitivity, demons-
trated by the atto- and femtomolar detection of the
bioanalyte and confirmed by an AC analysis of the
binding molecules. Throughout the paper, these types
of experimental analyses are referred to as memristor
measurement mode. Although the high sensitivity is
proven, the technique shows a limitation marked by
the initial conditions of the gap and the charge sign
of the analyte to be detected. The charge sign of the
analyte (positive or negative) governs the processes
of the opening or closing of the VGAP. When the initial
measurements do not lead to the creation of a gap
(VGAP = 0), the technique is limited to detect analytes of
the charge type that will open it. Therefore, detection
of analytes of the opposite charge, e.g., analytes with
a gap-closing effect, will remain challenging. The
solution to this fundamental problem lies in establishing
an external control of the initial state (gap opening)
conditions.
Finally, the previously reported memristor biosensors
operate in dry conditions, which limits the range of
their potential applications in, for example, clinical
or point of care (POC) diagnostics, where the assays
are dominantly performed in liquid phase. Such dry
measurement formats aim to overcome the Debye
length limitation that decreases the screening length
capacity of these biosensors owing to the presence of
other ions in the sample [26]. Note that humidity
control is, however, crucial for the dry format, because
minor humidity perturbations would affect the width
of the gap [40]. In the context of POC applications,
typically applied out of the laboratory environment
with a minimum technological setup, measurements
in liquid phase are preferable. Very recently, such
measurements were demonstrated using a bare device
for pH sensing of buffer solutions [41], but biosensing
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
3 Nano Res.
in liquid samples is still unexplored.
Here, we attempt to solve the number of aforemen-
tioned shortcomings in the existing sensors measured
in memristor mode and present the first liquid
integrated bio-memristor based on a large array of
silicon-nanowire-based electrical devices. The pre-
sence of ionic species attached in proximity to the
semiconducting channel introduces a capacitive effect
that violates the memristor characteristics by modifying
the VGAP. The width change of this gap can be correlated
to the concentration of molecules in the sample. We
further engineered the gap opening process by involving
the third electrode, which acts as a gap-control
terminal (gate voltage, VG) that mimics the presence
of the surrounding charged species of the desired sign
(Figs. 1(a) and 1(b)). With the help of this electrode,
the creation of the VGAP in the output characteristics
can be done on purpose, to assure detection of analytes
of any charge sign at low concentrations. Finally, we
compared the bio-memristor performance, in terms
of its detection range and sensitivity, to that of the
already-known FET mode by operating the same
device. The demonstration is done for detection of
the Ebola VP40 matrix protein, which represents high
emerging relevance owing to its recent severe outbreaks
and high fatality rate [42]. A portable, miniaturized
biosensor that can be used in a non-specialized
laboratory or location is urgently needed for the low-
cost early detection of the disease, which would help
to prevent its spread. The most common diagnostic
methods for Ebola are the polymerase chain reaction
(PCR) and the enzyme-linked immunosorbent assay
(ELISA) [43]. To date, few alternative approaches have
been proposed, such as plasmonic [44] or surface
acoustic wave immunosensors [45] and fluorescence
DNA sensors based on nanoporous membrane systems
[46]. Despite the good results achieved, a purely
electrical biosensor would be preferable for the
Figure 1 Concept of the present work: (a) VP40 proteins from an Ebola virus bind to antibodies immobilized on a HC-FET and (b) theviolation of the memristor zero-crossing signature of the biosensor is measured. The opening or closing of the voltage minimum gap(VGAP = V1 – V0) when sweeping current to source-to-drain voltage VSD is plotted. (c) Device used in this work. An array of transistors isplaced surrounding a common reference electrode. The electrode design of the area marked in the white window is shown in (d).(e) Magnification of the same area on a fabricated device. S = source electrode, D = drain electrode, R = reference electrode. The insetin (e) presents a SEM image of the nanowire pattern. (f) Transfer characteristic of the HC-FET at fixed VSD = 0.1 V in dry conditions.Gate voltage VG is applied via a back gate. Inset shows output characteristics at three different VG. Dotted lines in both graphs representthe gate-to-drain current (leakage). (g) Transfer characteristic of the HC-FET at fixed VSD = 0.1 V in liquid conditions (deionized water).Upper inset shows output characteristics at three different VG. Dotted lines in both graphs represent the gate-to-drain current (leakage).Lower inset presents a TEM image of the nanowires’ cross-section.
| www.editorialmanager.com/nare/default.asp
4 Nano Res.
abovementioned reasons, leading to the implementation
of several optical setup-free nanosensors in a tiny chip.
We propose a direct, label-free, and rapid method for
the detection of VP40—a protein that contributes 40%
to the Ebola virus protein mass [47], making it the
ideal candidate as a biomarker for diagnosis. The effects
of the binding of this protein on surface-immobilized
antibodies were detected using the previously reported
high-performance honeycomb patterned nanowire FETs
(HC-FETs) [48, 49], showing remarkable sensitivity.
The gap-control terminal was used during the bio-
sensing, which, unlike in other memristor biosensors
[37, 38, 40], will allow us to set the desired initial
conditions for the detection of proteins of any charge
sign.
2 Results and discussion
2.1 Engineering of the voltage gap opening
A chip with an array of HC-FETs with a common
reference electrode was used in this study (Figs. 1(c)
and 1(d)); nanowires on this electrode were patterned
as a honeycomb structure (Fig. 1(e)) owing to their
proven higher sensitivity compared to that of linear
structures [49]. FETs with 32.5-nm-wide (average)
nanowire channels, lightly doped (1014 cm−2) with
phosphorus and patterned in a honeycomb shape
by electron beam lithography (EBL), were used for all
experiments (Fig. 1(e)). Source and drain regions were
highly doped with phosphorus. Source and drain
electrodes, as well as the VGAP-controlling reference
electrode for liquid gating, were fabricated via
deposition of stacked Ti and Ag layers and lift-off
processing. The devices were insulated with a 2-μm-
thick layer of SU-8 resist, leaving the nanowire area
and the reference electrode exposed to the environment.
The use of the EBL technique ensured good sensor-
to-sensor reproducibility on a wafer-scale fabrication.
More details on fabrication can be found in the
experimental section. The fabrication process resulted
in n-type FETs with a working range below VG = 30 V
in dry conditions (Fig. 1(f)) using the back gate,
whereas the VG needed for liquid conditions (Fig. 1(g))
using the reference electrode decreased by 10 times
owing to the thinner top oxide layer. Electrical mea-
surements were performed using a tip probe station
with PH150 micropositioners (SÜSS MicroTec, Garching
bei München, Germany) to connect the contact pads
and a source meter unit (2604B source meter unit,
Keithley Instruments GmbH, Germering, Germany).
The VGAP between the minimum current peaks in the
forward and backward sweep of VSD was calculated
using the following equation (Fig. 1(b))
VGAP = V1 – V0 (1)
where V1 and V0 are the current-minima voltages for the
two sweeping directions. For all the measurements,
the fastest scan rate allowed by the data acquisition
software (Matlab, Mathworks, Natick, MA, USA) was
applied (t = 10 s per sweep). As shown in Fig. S1 in the
Electronic Supplementary Material (ESM), at slower
speeds VG values that were more negative were needed
to open the VGAP, owing to easier movement of the
charge carriers and faster recovery of the chip. A fast
scan rate facilitated obtaining the gap at lower voltages
and acquiring the results in a shorter time.
To our knowledge, we demonstrate for the first
time the possibility to achieve and control the device
response in memristor measurement mode, using
external guidance: back gate for measurements in dry
conditions and microfabricated top gate for liquid
integrated experiments. The gating mimics the presence
of surrounding charged molecules, allowing the
opening or closing of the VGAP in a controlled manner to
set the initial conditions as needed. First, we performed
a dry conditions test using the back gate control, by
sweeping VSD between −2.5 and 2.5 V. Initially, no gap
opening was observed for zero or positive VG values
(Fig. 2(a)). Positive charges attract the negative carriers
at the channel region, facilitating their movement and,
therefore, the recovery of the conditions along the
sweep without any memory effect. Both current minima
crossed each other through the zero voltage value,
resulting in VGAP = 0 V. In contrast, a negative increase
of VG initiated the appearance of the characteristic
hysteresis in the I–VSD curve (Figs. 2(b) and 2(c)). The
separation between both peaks, i.e., the VGAP, reveals
an increasing tendency with VG increase. Once the gap
started to appear, VGAP dramatically increased with
small VG changes starting at approximately VG = −3 V,
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
5 Nano Res.
Figure 2 Voltage minima gap VGAP control with back gate in dry
conditions. (a)–(c) Drain-current–source-to-drain voltage I–VSD
curves at gate voltage VG from 0 to −3 V. As the VG decreases, the
VGAP widens. The current at the positive branch of the VSD sweep
also decreases, making the recovery more difficult, which results
in low currents at the negative branch of the backward sweep.
The orange region indicates VGAP. (d) Calibration of VGAP as a
function of VG for dry conditions. The gray area indicates the
most sensitive region, where the VGAP changes most dramatically
with the VG. The inset shows a cross-sectional diagram of the
device, where VG was applied via a back gate. (e)–(g) I–VSD
curves at various VG values in liquid conditions. (h) Calibration
of VGAP as a function of VG for liquid conditions (deionized
water). The gray area indicates the most sensitive region. The
inset shows a cross-sectional diagram of the device, where the VG was applied via the reference electrode in contact with the liquid.
as follows (see the grey region in Fig. 2(d))
GAP
G
4.5V
V
(2)
indicating a high sensitivity to changes in the surroun-
ding environment, as summarized in Fig. 2(d). Note
that the current at the positive branch of the sweep
decreased as the VG value became more negative,
owing to a decrease in the channel conductivity near
the source region. A large VG drop reduced more
electron carriers within the channel near the source
region. At the most negative VG value, the current was
low, and the device did not have enough time to
recover; this reflected as low current values even in
the negative branch of the sweep.
Next, the possibility to control the gap when a
liquid sample and the reference electrode were used
was tested by depositing a 100-μL deionized water
droplet. Liquid measurements confirmed the earlier
demonstrated trend in the gap opening as a function
of VG increase (Figs. 2(e)–2(h)), with the only difference
being that 10-fold lower VG were necessary, which was
attributed to the thin top oxide layer (5-nm thickness)
compared to the back oxide (200-nm thickness). By
defining the gate potential through the liquid
environment, the gap started to be seen at −175 mV,
with an abrupt hysteresis increase of
GAP
G
60V
V
(3)
in the most sensitive range (13-fold more sensitive
than when using a back gate, as indicated by the grey
region in Fig. 2(h)).
2.2 Bare memristor and transistor performance for
pH sensing
Considering that the application of a VG is analogous
to the presence of charged species of a certain sign,
the change in the charge content in a liquid solution
close to the nanowires’ surface should affect the VG
required to open the gap. Positive charges would
require the application of a more negative VG to com-
pensate their opposite effect, and vice versa. Phosphate
buffered saline (PBS) drops (100 μL), 100-fold diluted
and adjusted to various pH values, were deposited on
the sensors, and the VGAP was calculated at different VG.
As observed in Fig. 3(a), as the PBS solution became
more basic (i.e., less positive charges/protons), the VG
needed to be increased to more positive values to open
the gap. For acidic samples, the VG needs to be increased
by 50 mV/pH to maintain a VGAP value of 1.5 V
| www.editorialmanager.com/nare/default.asp
6 Nano Res.
(Fig. 3(b)), while the sensitivity reduces to 7 mV/pH
in the basic range.
A comparative test using the sensor in the traditional
FET mode (Figs. 3(c) and 3(d)) shows the expected
behavior, with a shift of the transfer curve to the
positive direction as the pH increases. By measuring
in the subthreshold regime (at I = 0.01 μA), a sensitivity
of 35 mV/pH was obtained in the most sensitive acidic
range, which slightly decreased, as expected, for the
basic range (7 mV/pH). The larger amount of mV/pH
needed to maintain VGAP = 1.5 V in the case of the
memristor mode could be a sign of its higher sensitivity
compared to that of the FET mode. However, the loss
of linearity for the smaller VGAP values and the crossing
of the different curves indicate that the control of this
mode requires further studies to improve it.
2.3 Specific Ebola detection using bio-memristor
Finally, we applied the memristor mode for the
specific detection of the Ebola VP40 matrix protein
in a liquid phase and compared it with that using
the FET operation mode. Specific antibodies against
the VP40 protein were immobilized onto the amino
modified nanowires’ surface, using carbodiimide
chemistry. The steps followed are shown in Fig. 4(a)
and explained in more detail in the Experimental
section. Briefly, the bare nanowires were plasma
activated to generate hydroxyl groups, onto which
3-aminopropyl triethoxysilane (APTES) was attached
by incubating for 1 h in an ethanolic solution. After
cleaning the surface with ethanol and drying the
devices for 30 min at 120 °C, the antibodies were
immobilized on the APTES-modified surface by
incubation for 1 h in 1× PBS in the presence of 1-ethyl-
3-(3-dimethylaminopropyl)carbodiimide hydrochloride
(EDC) and N-hydroxysuccinimide (NHS), followed
by a final rinse with PBS and a blocking step of
the remaining free surface sites with bovine serum
albumin (BSA) (0.5 mg/mL). More details on the
biofunctionalization can be found in the experimental
section.
For the analysis of the biosensing response, PBS
droplets (20 μL) with increasing VP40 concentrations
were incubated for 30 min on the biosensors, and then
cleaned with PBS to remove loosely bound proteins.
After rinsing with sodium phosphate buffer (SP) at a
5-mM concentration, a 100-μL droplet of the SP buffer
was deposited to obtain an increased Debye length
and thus higher sensitivity during the measurements.
The signals using both memristor and FET approaches
Figure 3 Calibration of the sensor sensitivity toward different pH solutions. (a) and (b) Using the memristor approach; (c) and (d)using the field-effect transistor approach.
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
7 Nano Res.
were acquired and compared.
Figures 4(b) and 4(c) show the response of the
biosensor following the memristor approach. A shift
of the VGAP–VG curve to the right could be observed with
increasing antigen concentration, indicating molecule
detection at femtomolar levels, with 6 fM being the
smallest detected concentration. The detection range
and the smallest detected concentration was the same
when using the FET measuring mode, although the
signal change was more pronounced in the latter case
(Figs. 4(d) and 4(e)), with a shift of the transfer curve
toward higher VG values. A specificity test was done
using staphylococcal enterotoxin B (SEB). Toxins pro-
duced by this bacterium cause similar initial symptoms
to those of Ebola (i.e., fever, diarrhea, vomiting) [50].
The results, depicted in Fig. 4(e), showed that the
deviation of the VG caused by nonspecific binding
was significantly lower.
These results agree with the detection capabilities
of the honeycomb patterned FETs, as published
previously [51], and improve the picomolar detection
level expected for rapid Ebola diagnosis in point-of-care
situations [52]. Furthermore, the detection range is
similar to that of another recent Ebola nanosensor,
where detection was possible at 50 fM, although
still requiring labeling for optical detection [46]. A
comparison of different Ebola detection techniques is
summarized in Table 1.
As a reference measurement and for comparison,
the VP40 assay was performed using the standard
ELISA format to confirm the immunorecognition of
the antigen by the antibody and to compare the results
of our devices with those of a traditional and well-
established technique that relies on optical transduction.
The antibody-antigen recognition was carried out using
horseradish peroxidase (HRP)-modified antibodies.
Figure 4 Sensor surface functionalization and results of the VP40 detection using such surfaces. (a) Functionalization steps: (i) bare honeycomb nanowires, (ii) silanization with APTES, (iii) antibody immobilization, and (iv) detection of the VP40 protein. (b) and (c)Calibration of the biosensor by VP40 detection in memristor mode. The inset in (c) shows the optical detection of VP40 by the enzyme-linked immunosorbent assay technique using horseradish peroxidase-modified antibodies. (d) and (e) Calibration of the biosensor by VP40detection in field-effect transistor mode, including the specificity test with the SEB (black data points).
| www.editorialmanager.com/nare/default.asp
8 Nano Res.
Incubation for 10 min in tetramethylbenzidine (TMB)
resulted in the development of a blue color that turned
to yellow upon addition of a stop solution (0.2 M
sulfuric acid). The absorbance was obtained at 450 nm
using an ELISA multiwell reader. The result is shown
in the inset of Fig. 4(e). Measurable signals could be
obtained from the antibody–antigen binding events,
with a linear increase of the absorbance as the amount
of antigen increased up to 50 nM, and slowly saturating
afterwards. The smallest concentration that could
be sensed accurately was 6.25 nM. The two electrical
readout modes, memristor and FET, outperformed
the traditional ELISA technique in both detection
capabilities (by six orders of magnitude) and
miniaturization.
3 Conclusions
A gate-controlled bio-memristor based on a large
nanowire area patterned in a honeycomb structure
and capable of sensing biomolecules directly in liquid
environments is presented here for the first time. The
gate control is used to engineer the opening of the VGAP
during the measurements, allowing initial conditions
that permit the detection of analytes of any charge
sign to be set. The biosensing capability of this device
at the femtomolar level is shown for the Ebola VP40
matrix protein, which is better than the results of
both the traditional method (ELISA) and other recent
nanosensors [46]. The biosensor can perform highly
sensitive, label-free measurements in two ways: as
an ion-sensitive FET and through a violation of its
memristor zero-crossing signature by opening a VGAP.
The memristor approach—a recent method—has
been demonstrated directly on liquid samples, which
would allow its application in a more simplified
setup that does not require humidity control. The
controllability of the VGAP using external guidance
through the gate electrode has been demonstrated on
a honeycomb nanowire pattern, rather than on a
single or a few linear nanowires [37, 38]. Furthermore,
the engineering of the VGAP opening (the basic property
used for the quantification) has been demonstrated in
dry (through the back gate) and liquid (through the
liquid gate, using the reference electrode) conditions.
The measurements can then be performed either by
measuring the VGAP change at a fixed VG or by measuring
the changes in the VG needed to maintain a fixed VGAP.
We predict that this device can be of great importance
for improving the control of memristor biosensors
in the ultrasensitive detection of both positively and
negatively charged molecules in realistic point-of-care
situations and for the rapid diagnosis of fast-spreading
diseases.
Table 1 Comparison of different Ebola detection techniques
Technique Target Lowest detection Miniaturization Particular drawbacks
Optical (plasmonic) [44] Live virus 106 pfu/mL + Special safety facilities
Surface acoustic wave [45] Whole virus 104 pfu/mL +
Fluorescence (flow cytometry) [53] Whole virus (DNA staining)
105 pfu/mL – – Fluorescent labeling required
Plaque assay [53] Live virus 101 pfu/mL – – Special safety facilities, several days duration
Quantitative reverse transcription-polymerase chain
reaction (qRT-PCR) [53]
RNA 103 pfu/mL – – Special training and equipment, hours duration
TEM [53] Virus particles (VP) 106 VP/mL – – Special equipment, hours duration
Optical (nanoparticle based luminescence) [46]
DNA fM – – Nanoparticle labeling required
ELISA (optical, absorbance) (inset Fig. 4(e))
VP40 nM – – Labeled antibody for colorimetric enzymatic reaction required
This work (memristor/FET) VP40 fM + +
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
9 Nano Res.
4 Experimental
4.1 Device fabrication
Devices were fabricated on silicon-on-insulator wafers
(p-type, ~10 Ω·cm) with 200-nm-thick oxide and
50-nm-thick top silicon layers. Phosphorus at a concen-
tration of 1014 cm−2 was implanted in the top layer.
The dopant was activated by rapid thermal annealing
at 1,000 °C. The active region was formed and isolated
using inductively coupled plasma reactive ion etching
(ICP-RIE). Then, a heavy implantation of phosphorus
was placed into the source and drain region, followed
by dopant activation by rapid thermal annealing
at 1,000 °C. The honeycomb pattern of 50-nm-wide
nanowires was defined by electron beam lithography
and ICP-RIE. The total sensing surface area of the
nanowires was 365.2 μm2. A gate oxide layer of 5 nm
was grown using a thermal furnace at 900 °C. A
transmission electron microscope (TEM) observation
revealed that the etching and oxide growth processes
resulted in final nanowire widths of 40 nm at the
bottom and 25 nm at the top (see bottom inset in
Fig. 1(g)). Subsequently, stacked 50-nm Ti and 200-nm
Ag layers were evaporated to form the reference
electrode as well as the external pads for source and
drain. The purpose of the reference electrode was to
define the potential of the liquid samples. It was used
as the top liquid gate for the measurements in liquid
conditions.
The devices were isolated by spin coating a 2-μm
SU-8 resist. The active region (including the nanowire
channel region and the reference electrode) were left
exposed to the samples. The structure of the reference
electrode was completed by an electrochemical reaction
in 0.1 M KCl. A droplet of the buffer was deposited
on the electrode, and 1 V bias was applied from the
electrode in the presence of a grounded platinum
wire. The reaction created an Ag/AgCl layer on top of
the Ag electrode.
As a result of the fabrication process, n-type HC-FETs
(Fig. 1(e)) were obtained, with a nanowire region that
was 10-μm long and 150 μm away from the reference
electrode, as described previously [49].
4.2 Silicon nanowire biofunctionalization
A widely used biofunctionalization procedure was
followed for the immobilization of the antibodies on
the nanowires’ surface. As shown in Fig. 4(a), first,
air-plasma activation was applied to increase the
hydroxyl group content, as anchoring points for the
molecules immobilized in the next step. Then, a
silanization process was carried out using APTES.
The plasma-activated devices were immersed for 1 h
in an ethanolic solution containing 2.5% APTES and
5% deionized water. After rinsing with ethanol, they
were dried in an oven at 120 °C. The antibodies were
then directly immobilized onto the reactive amino
groups of APTES by incubating them in 0.1 mg/mL
PBS for 1 h in the presence of 10 mM EDC and 5 mM
NHS to allow activation of the carboxylic groups in the
antibodies. The unbound antibodies were removed
by rinsing with PBS, and the remaining free surface
was blocked against nonspecific adsorption by
incubation in 0.5 mg/mL BSA in PBS. After a final rinse,
the devices were ready and kept in PBS at 4 °C until
their use for the biosensing experiment.
4.3 ELISA procedure for VP40 detection
First, antibodies were labeled with HRP, using the
Lightning-Link HRP conjugation kit (Innova Biosciences,
Cambridge, UK). This kit enables the modification of
antibodies by using a modifier for direct labeling
with HRP, with the further addition of a quencher to
stop the reaction and obtain the conjugated antibodies
without the need of any separations.
Then, the ELISA test was conducted. For this test,
different antigen concentrations were adsorbed in
a multiwall plate by incubation in a carbonate/
bicarbonate buffer, pH 9.3, for 1 h. After rinsing to
remove the loosely bound antigen molecules, the
unoccupied areas of the plate wells were blocked by
incubation in a PBS solution with 2% BSA and 0.05%
Tween 20 for 1 h. Then, the wells were rinsed again
and incubated with antigen-specific, HRP-labeled
antibodies in PBS. After a final rinse to remove unbound
antibodies, the TMB color substrate was added for
10 min. HRP oxidized the TMB, resulting in a blue
color, which changed to yellow after the addition of
0.2 M sulfuric acid as a stop solution. The absorbance
at 450 nm, which depended on the concentration of
bound antibodies, and, therefore, of adsorbed antigens,
| www.editorialmanager.com/nare/default.asp
10 Nano Res.
was measured using a Tecan microplate reader (Tecan
Group Ltd., Männedorf, Switzerland).
Acknowledgements
This work was financed via the German Research
Foundation (DFG) within the Cluster of Excellence
“Center for Advancing Electronics Dresden (CfAED)
EXC 1056” and the “ICT Consilience Creative Program”
(No. IITP-R0346-16-1007) supervised by the Institute
for Information and Communications Technology
Promotion (IITP), Republic of Korea.
Electronic Supplementary Material: Supplementary
material (optimization of the drain-to-source voltage
sweeping rate) is available in the online version of
this article at https://doi.org/10.1007/s12274-017-1720-2.
References
[1] Liu, J.; Xie, C.; Dai, X. C.; Jin, L. H.; Zhou, W.; Lieber,
C. M. Multifunctional three-dimensional macroporous
nanoelectronic networks for smart materials. Proc. Natl.
Acad. Sci. USA 2013, 110, 6694–6699.
[2] Zörgiebel, F. M.; Pregl, S.; Römhildt, L.; Opitz, J.; Weber,
W. M.; Mikolajick, T.; Baraban, L.; Cuniberti, G. Schottky
barrier-based silicon nanowire pH sensor with live sensitivity
control. Nano Res. 2014, 7, 263–271.
[3] Gao, X. P. A.; Zheng, G. F.; Lieber, C. M. Subthreshold regime
has the optimal sensitivity for nanowire FET biosensors.
Nano Lett. 2010, 10, 547–552.
[4] Cui, Y.; Wei, Q. Q.; Park, H.; Lieber, C. M. Nanowire
nanosensors for highly sensitive and selective detection
of biological and chemical species. Science 2001, 293,
1289–1292.
[5] Haes, A. J.; Van Duyne, R. P. A nanoscale optical biosensor:
Sensitivity and selectivity of an approach based on the
localized surface plasmon resonance spectroscopy of triangular
silver nanoparticles. J. Am. Chem. Soc. 2002, 124, 10596–
10604.
[6] Vu, X. T.; GhoshMoulick, R.; Eschermann, J. F.; Stockmann,
R.; Offenhäusser, A.; Ingebrandt, S. Fabrication and appli-
cation of silicon nanowire transistor arrays for biomolecular
detection. Sen. Actuators, B Chem. 2010, 144, 354–360.
[7] Patolsky, F.; Zheng, G. F.; Lieber, C. M. Fabrication of silicon
nanowire devices for ultrasensitive, label-free, real-time
detection of biological and chemical species. Nat. Protoc.
2006, 1, 1711–1724.
[8] Patolsky, F.; Zheng, G. F.; Hayden, O.; Lakadamyali, M.;
Zhuang, X. W.; Lieber, C. M. Electrical detection of single
viruses. Proc. Natl. Acad. Sci. USA 2004, 101, 14017–14022.
[9] Patolsky, F.; Timko, B. P.; Yu, G. H.; Fang, Y.; Greytak, A. B.;
Zheng, G. F.; Lieber, C. M. Detection, stimulation, and
inhibition of neuronal signals with high-density nanowire
transistor arrays. Science 2006, 313, 1100–1104.
[10] Daniels, J. S.; Pourmand, N. Label-free impedance biosensors:
Opportunities and challenges. Electroanalysis 2007, 19,
1239–1257.
[11] Sharma, R.; Deacon, S. E.; Nowak, D.; George, S. E.;
Szymonik, M. P.; Tang, A. A. S.; Tomlinson, D. C.; Davies,
A. G.; McPherson, M. J.; Wälti, C. Label-free electrochemical
impedance biosensor to detect human interleukin-8 in serum
with sub-pg/mL sensitivity. Biosens. Bioelectron. 2016, 80,
607–613.
[12] Lin, Z. Y.; Chen, L. F.; Zhang, G. Y.; Liu, Q. D.; Qiu, B.;
Cai, Z. W.; Chen, G. N. Label-free aptamer-based electro-
chemical impedance biosensor for 17β-estradiol. Analyst
2012, 137, 819–822.
[13] Medina-Sánchez, M.; Ibarlucea, B.; Pérez, N.; Karnaushenko,
D. D.; Weiz, S. M.; Baraban, L.; Cuniberti, G.; Schmidt, O.
G. High-performance three-dimensional tubular nanomembrane
sensor for DNA detection. Nano Lett. 2016, 16, 4288–4296.
[14] Chen, K. I.; Li, B. R.; Chen, Y. T. Silicon nanowire field-
effect transistor-based biosensors for biomedical diagnosis
and cellular recording investigation. Nano Today 2011, 6,
131–154.
[15] Liu, S.; Guo, X. F. Carbon nanomaterials field-effect-
transistor-based biosensors. NPG Asia Mater. 2012, 4, e23.
[16] Schütt, J.; Ibarlucea, B.; Illing, R.; Zörgiebel, F.; Pregl, S.;
Nozaki, D.; Weber, W. M.; Mikolajick, T.; Baraban, L.;
Cuniberti, G. Compact nanowire sensors probe microdroplets.
Nano Lett. 2016, 16, 4991–5000.
[17] Karnaushenko, D.; Ibarlucea, B.; Lee, S.; Lin, G.; Baraban, L.;
Pregl, S.; Melzer, M.; Makarov, D.; Weber, W. M.; Mikolajick,
T. et al. Light weight and flexible high-performance diagnostic
platform. Adv. Healthc. Mater. 2015, 4, 1517–1525.
[18] Yang, Y. B.; Yang, X. D.; Zou, X. M.; Wu, S. T.; Wan, D.;
Cao, A. Y.; Liao, L.; Yuan, Q.; Duan, X. F. Ultrafine graphene
nanomesh with large on/off ratio for high-performance
flexible biosensors. Adv. Funct. Mater. 2017, 27, 1604096.
[19] Yang, Y. B.; Yang, X. D.; Tan, Y. N.; Yuan, Q. Recent
progress in flexible and wearable bio-electronics based on
nanomaterials. Nano Res. 2017, 10, 1560–1583.
[20] Bergveld, P. Development of an ion-sensitive solid-state
device for neurophysiological measurements. IEEE Trans.
Biomed. Eng. 1970, 17, 70–71.
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
11 Nano Res.
[21] Pregl, S.; Heinzig, A.; Baraban, L.; Cuniberti, G.; Mikolajick,
T.; Weber, W. M. Printable parallel arrays of Si nanowire
Schottky-barrier-FETs with tunable polarity for comple-
mentary logic. IEEE Trans. Nanotechnol. 2016, 15, 549–556.
[22] Baraban, L.; Zörgiebel, F.; Pahlke, C.; Baek, E.; Römhildt, L.;
Cuniberti, G. Lab on a wire: Application of silicon nanowires
for nanoscience and biotechnology. In Nanowire Field Effect
Transistors: Principles and Applications; Kim, D. M.; Jeong,
Y. H., Eds.; Springer: New York, 2014; pp 241–278.
[23] Pregl, S.; Weberk W. M.; Nozaki, D.; Kunstmann, J.;
Baraban, B.; Opitz, J.; Mikolajick, T.; Cuniberti, G. Parallel
arrays of Schottky barrier nanowire field effect transistors:
Nanoscopic effects for macroscopic current output. Nano
Res. 2013, 6, 381–388.
[24] Namdari, P.; Daraee, H.; Eatemadi, A. Recent advances in
silicon nanowire biosensors: Synthesis methods, properties,
and applications. Nanoscale Res. Lett. 2016, 11, 406.
[25] Shen, M. Y.; Li, B. R.; Li, Y. K. Silicon nanowire field-
effect-transistor based biosensors: From sensitive to ultra-
sensitive. Biosens. Bioelectron. 2014, 60, 101–111.
[26] Stern, E.; Wagner, R.; Sigworth, F. J.; Breaker, R.; Fahmy,
T. M.; Reed, M. A. Importance of the Debye screening length
on nanowire field effect transistor sensors. Nano Lett. 2007,
7, 3405–3409.
[27] Gao, N.; Gao, T.; Yang, X.; Dai, X. C.; Zhou, W.; Zhang,
A. Q.; Lieber, C. M. Specific detection of biomolecules in
physiological solutions using graphene transistor biosensors.
Proc. Natl. Acad. Sci. USA 2016, 113, 14633–14638.
[28] Presnova, G.; Presnov, D.; Krupenin, V; Grigorenko, V.;
Trifonov, A.; Andreeva, I.; Ignatenko, O.; Egorov, A.;
Rubtsova, M. Biosensor based on a silicon nanowire field-
effect transistor functionalized by gold nanoparticles for the
highly sensitive determination of prostate specific antigen.
Biosens. Bioelectron. 2017, 88, 283–289.
[29] Krivitsky, V.; Zverzhinetsky, M.; Patolsky, F. Antigen-
dissociation from antibody-modified nanotransistor sensor
arrays as a direct biomarker detection method in unprocessed
biosamples. Nano Lett. 2016, 16, 6272–6281.
[30] Ingebrandt, S. Bioelectronics: Sensing beyond the limit. Nat.
Nanotechnol. 2015, 10, 734–735.
[31] Laborde, C.; Pittino, F.; Verhoeven, H. A.; Lemay, S. G.;
Selmi, L.; Jongsma, M. A.; Widdershoven, F. P. Real-time
imaging of microparticles and living cells with CMOS
nanocapacitor arrays. Nat. Nanotechnol. 2015, 10, 791–795.
[32] Knopfmacher, O.; Tarasov, A.; Fu, W. Y.; Wipf, M.; Niesen,
B.; Calame, M.; Schönenberger, C. Nernst limit in dual-gated
Si-nanowire FET sensors. Nano Lett. 2010, 10, 2268–2274.
[33] Chua, L. O. Memristor—The missing circuit element. IEEE
Trans. Circuit Theory 1971, 18, 507–519.
[34] Strukov, D. B.; Snider, G. S.; Stewart, D. R.; Williams, R. S.
The missing memristor found. Nature 2008, 453, 80–83.
[35] Ascoli, A.; Slesazeck, S.; Mahne, H.; Tetzlaff, R.; Mikolajick,
T. Nonlinear dynamics of a locally-active memristor. IEEE
Trans. Circuits Syst. I Regul. Pap. 2015, 62, 1165–1174.
[36] Ascoli, A.; Tetzlaff, R.; Chua, L. O.; Strachan, J. P.; Williams,
R. S. History erase effect in a non-volatile memristor. IEEE
Trans. Circuits Syst. I Regul. Pap. 2016, 63, 389–400.
[37] Carrara, S.; Sacchetto, D.; Doucey, M. A.; Baj-Rossi, C.;
De Micheli, G.; Leblebici, Y. Memristive-biosensors: A new
detection method by using nanofabricated memristors. Sens.
Actuators B Chem. 2012, 171–172, 449–457.
[38] Tzouvadaki, I.; Jolly, P.; Lu, X. L.; Ingebrandt, S.; de
Micheli, G.; Estrela, P.; Carrara, S. Label-free ultrasensitive
memristive aptasensor. Nano Lett. 2016, 16, 4472–4476.
[39] Chua, L. If it’s pinched it’s a memristor. In Memristors and
Memristive Systems; Tetzlaff, R., Ed.; Springer: New York,
2014; pp 17–90.
[40] Puppo, F.; Dave, A.; Doucey, M. A.; Sacchetto, D.; Baj-Rossi,
C.; Leblebici, Y.; De Micheli, G.; Carrara, S. Memristive
biosensors under varying humidity conditions. IEEE Trans.
NanoBioscience 2014, 13, 19–30.
[41] Tzouvadaki, I.; Lu, X.; De Micheli, G.; Ingebrandt, S.;
Carrara, S. Nano-fabricated memristive biosensors for bio-
medical applications with liquid and dried samples. In 2016
38th Annual International Conference of the IEEE Engineering
in Medicine and Biology Society (EMBC), Orlando, USA,
2016, pp 295–298.
[42] Goodchild, S. A.; Dooley, H.; Schoepp, R. J.; Flajnik, M.;
Lonsdale, S. G. Isolation and characterisation of Ebolavirus-
specific recombinant antibody fragments from murine and
shark immune libraries. Mol. Immunol. 2011, 48, 2027–2037.
[43] Lucht, A.; Grunow, R.; Möller, P.; Feldmann, H.; Becker, S.
Development, characterization and use of monoclonal
VP40-antibodies for the detection of Ebola virus. J. Virol.
Methods 2003, 111, 21–28.
[44] Yanik, A. A.; Huang, M.; Kamohara, O.; Artar, A.; Geisbert,
T. M.; Connor, J. H.; Altug, H. An optofluidic nanoplasmonic
biosensor for direct detection of live viruses from biological
media. Nano Lett. 2010, 10, 4962–4969.
[45] Baca, J. T.; Severns, V.; Lovato, D.; Branch, D. W.; Larson,
R. S. Rapid detection of Ebola virus with a reagent-free,
point-of-care biosensor. Sensors 2015, 15, 8605–8614.
[46] Tsang, M. K.; Ye, W. W.; Wang, G. J.; Li, J. M.; Yang, M.;
Hao, J. H. Ultrasensitive detection of Ebola virus
oligonucleotide based on upconversion nanoprobe/nanoporous
membrane system. ACS Nano 2016, 10, 598–605.
[47] Elliott, L. H.; Kiley, M. P.; McCormick, J. B. Descriptive
analysis of Ebola virus proteins. Virology 1985, 147, 169–176.
| www.editorialmanager.com/nare/default.asp
12 Nano Res.
[48] Rim, T.; Kim, K.; Kim, S.; Baek, C. K.; Meyyappan, M.;
Jeong, Y. H.; Lee, J. S. Improved electrical characteristics
of honeycomb nanowire ISFETs. IEEE Electron Device
Lett. 2013, 34, 1059–1061.
[49] Rim, T.; Meyyappan, M.; Baek, C. K. Optimized operation of
silicon nanowire field effect transistor sensors. Nanotechnology
2014, 25, 505501.
[50] Marples, R. R.; Wieneke, A. A. Enterotoxins and toxic-shock
syndrome toxin-1 in non-enteric staphylococcal disease.
Epidemiol. Infect. 1993, 110, 477–488.
[51] Kim, K.; Park, C.; Kwon, D.; Kim, D.; Meyyappan, M.;
Jeon, S.; Lee, J. S. Silicon nanowire biosensors for detection
of cardiac troponin I (cTnI) with high sensitivity. Biosens.
Bioelectron. 2016, 77, 695–701.
[52] Kaushik, A.; Tiwari, S.; Dev Jayant, R.; Marty, A.; Nair, M.
Towards detection and diagnosis of Ebola virus disease at
point-of-care. Biosens. Bioelectron. 2016, 75, 254–272.
[53] Rossi, C. A.; Kearney, B. J.; Olschner, S. P.; Williams, P. L.;
Robinson, C. G.; Heinrich, M. L.; Zovanyi, A. M.; Ingram, M.
F.; Norwood, D. A.; Schoepp, R. J. Evaluation of ViroCyt®
virus counter for rapid filovirus quantitation. Viruses 2015,
7, 857–872.