379 J. Sensor Sci. & Tech. Vol. 26, No. 6, 2017
Journal of Sensor Science and Technology
Vol. 26, No. 6 (2017) pp. 379-385
http://dx.doi.org/10.5369/JSST.2017.26.6.379
pISSN 1225-5475/eISSN 2093-7563
Electrochemical Non-Enzymatic Glucose Sensor based on Hexagonal Boron Nitride
with Metal–Organic Framework Composite
Suresh Ranganethan1, Sang-Mae Lee
2, Jaewon Lee
3, and Seung-Cheol Chang
4+
Abstract
In this study, an amperometric non-enzymatic glucose sensor was developed on the surface of a glassy carbon electrode
by simply drop-casting the synthesized homogeneous suspension of hexagonal boron nitride (h-BN) nanosheets with a cop-
per metal−organic framework (Cu-MOF) composite. Comprehensive analytical methods, including field-emission scanning
electron microscopy (FE-SEM), Fourier-transform infrared spectroscopy (FT-IR), X-ray diffraction (XRD), cyclic vol-
tammetry, electrochemical impedance spectroscopy, and amperometry, were used to investigate the surface and elec-
trochemical characteristics of the h-BN–Cu-MOF composite. The FE-SEM, FT-IR, and XRD results showed that the h-BN–
Cu-MOF composite was formed successfully and exhibited a good porous structure. The electrochemical results showed a
sensor sensitivity of 18.1 µAµM−1cm−2 with a dynamic linearity range of 10−900 µM glucose and a detection limit of 5.5
µM glucose with a rapid turnaround time (less than 2 min). Additionally, the developed sensor exhibited satisfactory anti-
interference ability against dopamine, ascorbic acid, uric acid, urea, and nitrate, and thus, can be applied to the design and
development of non-enzymatic glucose sensors.
Keywords: Hexagonal boron nitride, copper metal–organic framework, non-enzymatic, electrochemical, glucose sensor
1. INTRODUCTION
Since the first introduction of a glucose biosensor by Clark in
1962 [1], numerous glucose sensors have been developed using
immobilized enzymes modified with various functional materials
[2]. However, the sensors developed with modified enzymes are
expensive and suffer from chloride ion poisoning and adsorption
of enzyme reaction intermediates [3]. To overcome these
limitations, during the last two decades, non-enzymatic biosensors
have been widely developed using various nanomaterials such as
metal nanoparticles, carbon nanotubes (CNTs), and graphene-
based nanomaterials [4,5], due to their excellent electron transfer
ability and significant resistance to chloride ion poisoning [6].
Besides the nanomaterials, metal–organic frameworks (MOFs)
have also received significant attention because their metal ions
exhibit a crystalline ordered structure, high porosity, large surface
area, thermal stability, and chemical tenability [7]. Particularly,
Cu-based MOFs (Cu-MOFs) have attracted significant attention in
biosensor research [8] and have been used in the determination of
biologically important compounds such as oxygen [9], H2O2 [10],
and dopamine [11]. Recently, a modified Cu-MOF has also been
used as a non-enzymatic glucose-sensing material [12]. In
addition, hexagonal boron nitride (h-BN), which is a wide-
bandgap semiconductor that replaces the C-C pair in the carbon
structure with an isoelectronic B-N bond, has been used in the
field of biosensors [13]. Importantly, h-BN is isostructural to
carbon-like materials such as graphite and CNTs, and thus, can be
used as a composite material to construct biosensors.
In the present study, an efficient method for synthesizing a
unique homogeneous suspension of few layered h-BN nanosheets
decorated with Cu-MOF (h-BN–Cu-MOF) was developed using a
simple sonication technique. The synthesized suspension was then
effectively modified onto a glassy carbon electrode (GCE) surface
to develop a non-enzymatic glucose sensor by using a simple
drop-casting method (Fig. 1). Therefore, the synergetic effect
created by the modified h-BN–Cu-MOF composite is expected to
improve the electro-oxidation sensitivity of glucose and enzyme-
mimic selectivity.
1Graduate Department of Chemical Materials, Pusan National University, 2
Busandaehak-ro, Geumjeong-gu, Busan 46241, Korea.
2Engineering Research Center for Net Shape and Die Manufacturing, Pusan
National University, 2 Busandaehak-ro, Geumjeong-gu, Busan 46241, Korea.
3College of Pharmacy, Molecular Inflammation Research Center for Aging
Intervention, Pusan National University, 2 Busandaehak-ro, Geumjeong-gu,
Busan 46241, Korea.
4Institute of BioPhysio Sensor Technology, Pusan National University, 2
Busandaehak-ro, Geumjeong-gu, Busan 46241, Korea.+Corresponding author: [email protected]
(Received: Sep. 21, 2017, Revised: Nov. 27, 2017, Accepted: Nov. 28, 2017)
This is an Open Access article distributed under the terms of the Creative
Commons Attribution Non-Commercial License(http://creativecommons.org/
licenses/bync/3.0) which permits unrestricted non-commercial use, distribution,
and reproduction in any medium, provided the original work is properly cited.
Suresh Ranganethan, Sang-Mae Lee, Jaewon Lee, and Seung-Cheol Chang
J. Sensor Sci. & Tech. Vol. 26, No. 6, 2017 380
2. EXPERIMENTAL
2.1 Materials and Chemicals
D-Glucose, ascorbic acid (AA), acetaminophen, uric acid (UA),
dopamine, NaOH, NaCl, copper nitrate trihydrate (Cu(NO3)2.3H2O),
1,3,5-benzenetricarboxylic acid (H3BTC), and dimethylformamide
(DMF) were purchased from Sigma-Aldrich (USA). h-BN mesh
powder (particle size: 1−5 μM) was purchased from Alpha Easer Co.
(Japan). All other reagents, of analytical grade, were purchased from
Sigma-Aldrich (USA) and used without further purification. All
aqueous solutions were prepared using deionized water (Milli-Q
water purifying system, 18 MΩ·cm−1).
2.2 Instrumentation
Cyclic voltammetry (CV) and amperometry were performed using
an electrochemical workstation (CompactStat, Ivium Technologies,
the Netherlands) with a conventional three-electrode cell system
consisting of a GCE (3 mm in diameter), a Ag/AgCl reference
electrode, and a platinum-wire auxiliary electrode. CV was carried out
in a NaOH solution by potential sweeping from +0.2 to +0.8 V, at a
scan rate of 50 mVs−1. For amperometry, the developed sensor was
inserted into an electrochemical cell, to which a 990 µL aliquot of
NaOH solution was added. The cell was then set and the sensor was
polarized at a potential of 0.60 V. After achieving a stable baseline
response with NaOH, a 10 µL aliquot of glucose sample was added,
and the current responses as a function of time were recorded. The
amperometry measurements were repeated to perform calibration and
investigate the sensor reproducibility and storage stability.
Electrochemical impedance spectroscopy (EIS) was performed
using an electrochemical workstation (VersaSTAT, Princeton Applied
Research, USA) in the frequency range of 100 KHz to 0.1 Hz at a DC
potential of 250 mV and AC potential of ±5 mV. To investigate the
surface characteristics, field-emission scanning electron microscopy
(FE-SEM), X-ray diffraction (XRD), and Fourier-transform infrared
spectroscopy (FT-IR) were carried out using the sensors modified
onto iridium tin oxide electrodes. FE-SEM was performed using a
field-emission scanning electron microscope (Hitachi S-4200, Japan)
operated at 15 kV, 150 W, and powder XRD patterns were collected
on an X-ray D/max-2200vpc (Rigaku Corporation, Japan) instrument
operated at 40 kV and 20 mA using Cu kα radiation (K = 0.15406).
2.3 h-BN–Cu-MOF-modified sensor preparation
The hydroxyl-functionalized aqueous dispersion of a few layers of
h-BN nanosheets was synthesized using the sonication-centrifugation
process [14]; 20 mg of the h-BN powder was dispersed in 10 mL of
deionized water and sonicated for 8 h, and the dispersion was
centrifuged at 3500 rpm for 12 min and filtered. The filtrate was then
collected as a “homogeneous” aqueous dispersion of a few layers of
h-BN nanosheets.
Cu-MOF was synthesized using a previously reported thermal
method [11, 15]; 0.55 g of Cu(NO3)2·3H2O was dissolved in 40 mL
of deionized water and mixed with 80 mL of 37.5 mM H3BTC
prepared in a 1:1 mixed solution of DMF and ethanol. The mixture
was kept in a water bath at 85°C for 24 h, and the blue powder of Cu-
MOF was separated by filtration and then dried at 80°C for 8 h. A
bare GCE was polished using 0.3-µm alumina slurries and sonicated
in ethanol for 10 min. After sonication, the GCE was rinsed
thoroughly with distilled water and dried at ambient temperature. To
prepare the h-BN–Cu-MOF composite, 3.0 mgmL−1 of the Cu-MOF
was dispersed in 1 mL of the h-BN dispersion under stirring for 10
min. A 5 μL aliquot of the composite suspension was drop-casted
onto the GCE surface and dried under ambient conditions.
Additionally, a 2 μL aliquot of Nafion solution (1.0 wt.% in ethanol)
was drop-cased to entrap the h-BN–Cu-MOF composite, which was
modified on the GCE. The constructed sensor is denoted as GCE–h-
BN–Cu-MOF/NF.
3. RESULTS AND DISCUSSIONS
3.1 Surface characteristics of h-BN–Cu-MOF composite
XRD was performed to study the structures of Cu-MOF without h-
BN and the h-BN–Cu-MOF composite (Fig. 2(A)). The XRD peaks
of Cu-MOF revealed its crystalline nature [16,17]; h-BN–Cu-MOF
spectra exhibited a diffraction peak at 26.31° for the (002) plane, and
at 42.93°, 44.16°, and 55.20° for the (100), (101), and (104) planes,
respectively. These spectra confirmed the formation of the h-BN–Cu-
MOF composite [18]. Fig. 2(B) shows the FT-IR spectra of Cu-MOF
and the h-BN–Cu-MOF composite. The FT-IR spectra of Cu-MOF
exhibit six prismatic crystals with Cu atoms, involving two carboxylic
Fig. 1. Illustration of proposed h-BN–Cu-MOF composite modified
non-enzymatic glucose sensor.
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Electrochemical Non-Enzymatic Glucose Sensor based on Hexagonal Boron Nitride with Metal–Organic Framework Composite
381 J. Sensor Sci. & Tech. Vol. 26, No. 6, 2017
“O” atoms, three water molecules, and a broad-band peak observed at
3432 cm-1, and conform to the O-H group in Cu-MOF [19,20].
However, the FT-IR spectra of the h-BN–Cu-MOF composite clearly
shows that the peak intensity of the O-H group and organic ligand
(BTC)2 C-O-C peak at 1110 cm-1 are completely diminished as
compared to case of pure Cu-MOF. The in-plane B-N transverse
stretching vibration peak at 1375 cm-1 and the out-of-plane B-N-B
bending vibration peak at 816 cm-1 can be suggested as a fingerprint
of sp2 bonds in h-BN sheets [21]. These results confirm that Cu-MOF
is effectively attached to h-BN sheets through the non-covalent
interaction induced by the sonication technique developed. The FE-
SEM image of pure Cu-MOF (Fig. 2(C)) shows crystals in micron-
sized tubular shape [22], while that of the h-BN–Cu-MOF composite
(Fig. 2(D)) shows a strong interaction of Cu-MOF with the h-BN
sheets. The Fe-SEM images clearly indicate that the h-BN sheets are
well-ordered in 2D form and are micron-sized.
3.2 Electrocatalytic oxidation of glucose
To investigate the electrocatalytic activity of the GCE–h-BN–Cu-
MOF/NF sensor, CV measurements were carried out in 0.15 M
NaOH at a scan rate of 50 mVs−1 with four prepared sensors: bare
GCE, GCE–h-BN/NF, GCE–Cu-MOF/NF, and GCE–h-BN–Cu-
MOF/NF. As shown in Fig. 3(A), in the absence of glucose, the
oxidation current in GCE–h-BN–Cu-MOF/NF is higher than that in
the other modified sensors, which can be attributed to the Cu (II)/Cu
(III) redox couple and the water-splitting process [23]. In the presence
of 0.3 mM glucose in 0.15 M NaOH, a clear glucose oxidation peak
appears at +0.60 V on GCE–h-BN–Cu-MOF/NF (inset in Fig. 3(A)).
However, the bare GCE, GCE–h-BN/NF, and GCE–Cu-MOF/NF
sensors show no detectable oxidation peaks of glucose. As illustrated
in Fig. 1, the electrocatalytic oxidation of glucose on GCE–h-BN–Cu-
MOF/NF underwent several steps: first, Cu-MOF was electrochemically
oxidized to Cu (III) species such as CuOOH [24], and then, glucose
was catalytically oxidized by Cu (III) species to form gluconic acid,
and at the same time, Cu (III) species were reduced to Cu (II) [25].
Fig. 2. (A) XRD images and (B) FT-IR spectra of the Cu-MOF and
the h-BN-Cu-MOF composite, (C) FE-SEM image of Cu-
MOF, and (D) FE-SEM image of the h-BN-Cu-MOF com-
posite.
Fig. 3. (A) Cyclic voltammograms of GCE-bare, GCE-h-BN/NF,
GCE-Cu-MOF/NF, and GCE-h-BN Cu-MOF/NF sensors in
the absence of glucose and in the presence of 0.3 mM glucose
(Inset). (B) EIS of GCE-bare, GCE-h-BN/NF, GCE-Cu-
MOF/NF, and GCE-h-BN-Cu-MOF/NF sensors in 0.1 M KCl
solution.
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Suresh Ranganethan, Sang-Mae Lee, Jaewon Lee, and Seung-Cheol Chang
J. Sensor Sci. & Tech. Vol. 26, No. 6, 2017 382
The four differently modified sensors were also employed for the
EIS study, as described in Section 2.2, and the Nyquist plot of each
sensor was obtained as shown in Fig. 3(B). It is well known that, in
EIS, the semicircle formed at a high-frequency region is related to the
charge transfer-limited process of the interface, while the linear
response formed at a low-frequency region is related to the diffusion-
limited process [26]. These results reveal that the oxidation peak at
+0.60 V shows a strong electrocatalytic activity of GCE–h-BN–Cu-
MOF/NF and a negligible charge transfer resistance of all sensors. It
is, therefore, suggested that the enhancement of electrocatalytic
activity is caused by the effective diffusion control process, the larger
size of the active surface area, and the high electron transfer rate of
the GCE–h-BN–Cu-MOF/NF sensor
In the mechanism of non-enzymatic glucose sensing, the
electrochemical behavior of Cu (II) and Cu (III) redox couple can be
considered as an essential factor for glucose oxidation at the sensor
surface [27,28]. The enhanced performance of the GCE–h-BN–Cu-
MOF/NF sensor is attributed to this factor through comparison with
other nanosized powders [29] and nanowires on film-type glucose
sensors [30]. Due to the large surface area of the nanoporous layers
of h-BN sheets decorated with Cu-MOF with high crystal quality, the
synergetic effect of the GCE–h-BN–Cu-MOF/NF sensor can explain
the selective glucose oxidation and enhancement of the charge
transfer resistance mechanism in Cu-MOF by π-π* transition with a
few layers of h-BN sheets. This also confirms that the electrons
generated are efficiently transferred from Cu-MOF to the sensor
surface with the high driving force created by the Schottky barrier
[31].
3.3 Sensor optimization
To achieve optimum sensor performance conditions, the operating
pH and potential were investigated by CV and amperometry
measurements toward glucose sensing. The non-enzymatic oxidation
of glucose produces gluconolactone and the simultaneous oxidation
continues to dehydrogenate, which is closely related to the presence
of hydroxide ions [26]. Thus, the hydroxide ions play an important
role in the oxidation of glucose at the GCE–h-BN–Cu-MOF/NF
sensor.
Fig. 4(A) shows the cyclic voltammograms of 0.3 mM glucose
prepared in deionized water with different concentrations of NaOH
(0−0.15 M). The oxidation peak current gradually increases with
NaOH concentration and slightly shifts to the negative-potential side
due to an easy oxidation of aldehyde and hydroxyl groups in glucose.
Accordingly, 0.15 M NaOH (pH 13.0) was used as the operating
medium in all experiments. A hydrodynamic voltammogram was
constructed using GCE–h-BN–Cu-MOF/NF with 0.1 mM glucose
(Fig. 4(B)). The maximum response was obtained at a potential of
+0.60 V and the signal response decreased dramatically beyond this
potential. Accordingly, +0.60 V was used as the operating potential in
all experiments.
3.4 Amperometric sensor calibration for glucose
Amperometry measurements were performed using the developed
sensors by adding glucose samples, as described in Sections 2.2 and
3.2. The current responses obtained using the bare GCE and GCE–h-
BN/NF sensors were negligible after the addition of glucose. As
shown in Fig. 5(A), in contrast, stable and immediate current
responses were observed after the addition of glucose and calibration
curves for glucose were constructed using the GCE–Cu-MOF/NF and
GCE–h-BN–Cu-MOF/NF sensors. As seen in Fig. 5(B), the current
responses reached steady-state values in less than 5 s. In the curve for
the GCE–Cu-MOF/NF sensor, the linear dynamic range was found to
be 10−900 μM and the sensitivity was calculated to be 11.0
μAμM−1cm−2. Responses to glucose using the GCE–h-BN–Cu-MOF/
NF sensor showed the same linearity ranges as high as 900 µM with
improved sensitivity of 18.1 μAμM−1cm−2, which was 1.5 times
higher than that of the GCE–Cu-MOF/NF sensor. The detection limit
of the GCE–h-BN–Cu-MOF/NF sensor was calculated to be 5.5 µM
and taken as six times the standard deviation of the current change
Fig. 4. (A) CVs of 0.3 mM glucose in different concentrations of
NaOH (0.0 M, 0.01 M, 0.03 M, 0.05 M, 0.08 M, 0.1 M, and
0.15 M) using the GCE-h-BN-Cu-MOF/NF sensor at a scan
rate of 50 mVs−1. (B) Effects of applied potential on the
amperometric response of the sensor to 0.1 mM glucose in
0.15 M NaOH (pH 13).
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Electrochemical Non-Enzymatic Glucose Sensor based on Hexagonal Boron Nitride with Metal–Organic Framework Composite
383 J. Sensor Sci. & Tech. Vol. 26, No. 6, 2017
due to the addition of a blank solution. The results obtained, therefore,
indicate that the h-BN–Cu-MOF composite-modified sensor exhibited
the highest sensitivity to glucose and improved the amperometric
sensor performance due to the synergetic effect of the h-BN–Cu-MOF
composite, as described in Section 3.2 [32].
3.5 Sensor performance characteristics
To the best of our knowledge, there are few reports on the non-
enzymatic glucose-sensing applications of Cu-MOF composites. In
this study, the performances of non-enzymatic electrochemical
sensors modified with Cu-MOFs were studied, and are summarized in
Table 1: Cu@C800: anthill-like Cu@carbon nanocomposites;
CuFe2O4: copper ferrite; MWCT: multiwall carbon nanotube; CuO:
Nanothorn Cu foam; Cu-MOF: copper metal–organic framework;
GCE: glassy carbon electrode. The summary clearly demonstrates
that the Cu-MOF-modified sensors rever good electrocatalytic ability
and can be used as non-enzymatic glucose sensors.
Interference is one of the major limitations in verifying the
selectivity of the developed non-enzymatic glucose sensor, because
the real samples may contain some co-existing biological compounds
in human blood serum. Fig. 6 shows the signal of the biosensors
following the addition of 0.1 mM DA, AA, UA, urea, and nitrate, and
0.2 mM NaCl. There was no significant change in the signal in
response to these compounds, indicating no substantial interference in
glucose detection. In addition, the use of Nafion film, as described in
Section 2.3, could avoid the interferences from AA and UA because
of its long backbone chain with negatively charged sulfonic groups
and their ionic properties. The interfering current responses for the
added compounds were only less than 6.0% that of glucose. The
relative standard deviation as a measure of inter-electrode
reproducibility was calculated to be 3.5%. This proved good
reproducibility of the developed sensor. Furthermore, only 4% loss of
the initial sensor sensitivity was observed after five-week storage of
the sensor at 4°C in dark.
Table 1. Sensors modified with Cu-MOF composites as non-enzymatic glucose sensing applications
Modified sensor Operating potential (V) Detection limit (μM) Dynamic Linear range (mM) Ref
Cu@C800 + 0.55 29.8 0.2−8.0 [33]
GCE/CuFe2O4MWCT +0.40 0.2 0.0005−1.4 [34]
CuO nanothorna +0.50 0.276 0.0002−2 [35]
Cu-MOF +0.70 1.0 0.005−2.8 [36]
GCE–h-BN–Cu-MOF/NF +0.60 5.5 0.01−0.90 This work
Fig. 5. (A) Amperometric current responses to glucose using the dif-
ferent sensors at a potential of +0.60 V in 0.15 M NaOH:
GCE-bare, GCE-h-BN/NF, GCE-Cu-MOF/NF and GCE-h-
BN-Cu-MOF/NF sensors. (B) Calibration curves for glucose
using the GCE-Cu-MOF/NF and GCE-h-BN-Cu-MOF/NF
sensors.
Fig. 6. Amperometric response to 0.1 mM glucose, 0.1 mM DA,
AA, UA, urea, NO3
- and 0.2 M NaCl at GCE-h-BN-Cu-MOF/
NF sensor in 0.15M NaOH.
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Suresh Ranganethan, Sang-Mae Lee, Jaewon Lee, and Seung-Cheol Chang
J. Sensor Sci. & Tech. Vol. 26, No. 6, 2017 384
4. CONCLUSIONS
A new method for synthesizing a non-enzymatic glucose sensor
was developed by simply drop-casting composite suspensions. The
composite was prepared based on a Cu-MOF composite decorated
with h-BN nanosheets, by a simple sonication technique developed.
The FE-SEM, FT-IR, and powder XRD results confirmed that the
composite was formed successfully with good porous structure. The
detection limit of the GCE–h-BN–Cu-MOF/NF sensor was 5.5 μM
glucose with a linear dynamic range of 10−900 µM. The developed
sensor exhibited enhanced features of sensitivity, reproducibility,
long-term stability, and anti-interference against electroactive species
in real biological samples such as DA, AA, and UA. These results
proved the usefulness of the sensor for glucose determination. Further
research is currently underway to develop a new non-enzymatic
glucose micro-sensor platform for real biological or clinical sample
analysis.
ACKNOWLEDGMENT
This work was supported by a 2-Year Research Grant of Pusan
National University.
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