Sensors and Actuators B 160 (2011) 287–294
Contents lists available at ScienceDirect
Sensors and Actuators B: Chemical
journal homepage: www.elsevier.com/locate/snb
Nanocomposite film based on graphene oxide for high performance flexible
glucose biosensor
Jian-Ding Qiu a,b,∗ , Jing Huang a , Ru-Ping Liang a
a
b
Institute for Advanced Study and Department of Chemistry, Nanchang University, Nanchang 330031, PR China
Department of Chemical Engineering, Pingxiang College, Pingxiang 337055, PR China
a r t i c l e
i n f o
Article history:
Received 31 May 2011
Received in revised form 19 July 2011
Accepted 24 July 2011
Available online 2 August 2011
Keywords:
Graphene oxide
Glucose oxidase
Chitosan–ferrocene
Biosensor
a b s t r a c t
A homogeneous chitosan–ferrocene/graphene oxide/glucose oxidase (CS–Fc/GO/GOx) nanocomposite
film was successfully constructed as a novel platform for the fabrication of glucose biosensor. The
morphologies and electrochemistry of the nanocomposite film were investigated by using scanning electron microscopy and electrochemical techniques including electrochemical impedance spectroscopy and
cyclic voltammetry, respectively. Results demonstrated that the uniformly dispersed GO within the CS
matrix could significantly improve the stability of GO and make it exhibit a positive charge, which was
more favorable for the further immobilization of biomolecules, such as GOx, with higher loading. Further attaching redox mediator ferrocene group (Fc) to CS could not only effectively prevent the leakage
of Fc from the matrix and retain its electrochemical activity, but also improve the electrical conductivity of CS and promote the electron-transfer between GOx and electrode. Biosensors based on this
CS–Fc/GO/GOx film had advantages of fast response, excellent reproducibility, high stability, and showed
a linear response to glucose in the concentration range from 0.02 to 6.78 mM with a detection limit
of 7.6 ␮M at a signal-to-noise ratio of 3 and exhibited a higher sensitivity of 10 ␮A mM−1 cm−2 . The
proposed strategy based on CS–Fc/GO nanocomposite for the immobilization of enzymes can be of practical relevance for the facile design of biosensors, as well as for the construction of new multifunctional
bioelectrochemical systems.
© 2011 Elsevier B.V. All rights reserved.
1. Introduction
Electrochemical biosensors based on nanomaterials, such as
carbon nanotubes [1–3], gold nanoparticles [4–6], metal oxides
[7,8], and semiconductors [9], have recently attracted considerable
attention in the medicine and food quality control field. In particular, because of the usefulness in diagnostic analysis of diabetes,
glucose biosensors based on carbon nanotubes and metal nanoparticles have been extensively studied in recent years [10–18].
Though the suitability of functional graphene for ultracapacitor and
energy storage applications has been studied extensively [19,20],
there are very few reports available that describe the suitability
of graphene for electrochemical biosensing applications [21–26].
Unlike rolled structured CNTs and metal nanoparticles with different sizes and shapes, graphene is a two-dimensional plenary
sheet with open structure, hence both sides of graphene could
be utilized for supporting enzymes. Therefore, it is expected to
∗ Corresponding author at: Institute for Advanced Study and Department of
Chemistry, Nanchang University, Nanchang 330031, PR China.
Tel.: +86 791 3969518.
E-mail addresses: [email protected], [email protected] (J.-D. Qiu).
0925-4005/$ – see front matter © 2011 Elsevier B.V. All rights reserved.
doi:10.1016/j.snb.2011.07.049
be a more promising enzymes carrier. However, the intrinsic van
der Waals interactions between layers of graphene easily results
in irreversible agglomeration or even restack to form graphite.
This problem has been encountered in all previous efforts aimed
at large-scale production of graphene through chemical conversion or thermal expansion/reduction [27–30]. The prevention of
aggregation is of particular importance for graphene because most
of their unique properties are only associated with individual
sheets [31]. This limitation can be overcome by the attachment of
other molecules or polymers onto the graphene surfaces. Zhang
and coworkers used deoxyribonucleic acid (DNA) for functionalizing graphene through ␲–␲ stacking interactions, the resulting
hydrophilic DNA/graphene nanocomposite ideally improved the
water-solubility of graphene and thus prevented their aggregation, meanwhile the obtained nanocomposite offered a promising
platform for immobilizing horseradish peroxidase for the development of novel electrochemical biosensors [32]. Saha et al. found that
the existence of polyacrylate on graphene and GO could assist the
dispersion of both in different aqueous buffer solutions with different pHs [33]. Recently, it has been reported that CS could form
a stable nanocomposite with GO through electrostatic attraction
and hydrogen bonding [34]. Owing to such efficient combination,
the resulting CS/GO nanocomposite is endowed with the excellent
288
J.-D. Qiu et al. / Sensors and Actuators B 160 (2011) 287–294
properties of the two independent components, such as the biocompatibility of CS and the outstanding electronic properties of
GO. Furthermore, the existence of CS on GO not only assists the
dispersion of GO in aqueous solution but also makes it exhibit a
positive charge, which is favorable for the further immobilization
of biomolecules through self-assembly.
As is well known, direct electron transfer between the active
site of the enzymes and the electrode surface is commonly forbidden, due to the fact that this redox center is commonly buried in
the globular structure of the protein. For this reason, the use of
electron transfer mediators has been widely employed during the
last few decades [35]. However, the operational simplicity desired
for biosensors is reduced by the use of these compounds in solutions. These drawbacks can be avoided by wiring the immobilized
enzyme in conducting or redox polymers [36,37]. In particular, the
modification of polymers with electron transfer mediators constitutes a simple, cheap and effective strategy for constructing
reagentless amperometric biosensors [38,39]. On the other hand,
enzymes are labile polypeptides that tend to be inactivated after
exposure to conditions commonly used for analytical determinations, and as a result, the stability and the sensitivity of the
electrode decay with time. To overcome this issue, the immobilization of enzymes in polysaccharide based supports has been widely
employed for the construction of biosensors [40,41]. Recently, we
and other groups reported the synthesis of biocompatible and conductive ferrocene branched chitosan (CS–Fc) [42,43], which was
used as redox polymer for the construction of reagentless enzyme
biosensors [44]. This novel redox active hybrid could not only effectively prevent the leakage of Fc from the matrix and improve the
electrical conductivity of CS but also show excellent biocompatibility for enzyme immobilization without destructing its native
structure [43,45].
In the present work, by taking the advantages of chitosan, GO
and ferrocene, we for the first time demonstrate the preparation of a novel platform for the fabrication of enzyme electrodes,
as illustrated in Scheme 1. As a common model enzyme, GOx
is negatively charged in physiological media, it can be immobilized on CS/GO through a self-assembly method to form a
multicomponent nanocomposite on a glassy carbon electrode. The
experimental results demonstrated that such an electrochemical platform not only effectively prevented the leakage of both
enzyme and mediator, preserved the native structure of the immobilized enzyme, but also provided an attractive route to promote
the electron transfer between GOx and electrode. Meanwhile, the
as-prepared CS–Fc/GO/GOx electrode exhibited excellent analytical performance toward the quantification of glucose, with a wide
linear range, excellent sensitivity, good reproducibility, and longterm stability. Therefore, this kind of redox active, conductive and
biocompatible nanocomposite material offers a promising platform for the development of novel electrochemical biosensors and
biomedical devices.
2. Experimental
2.1. Reagents
Graphite flake (99.8%, 325 mesh) was provided by Alfa Aesar
China Ltd. (China). Glucose oxidase (GOx, from Aspergillus niger,
EC 1.1.3.4. 150,000 units g−1 ), ␤-d(+)-glucose, chitosan (CS, 85%
deacetylation), sodium cyanoborohydride (NaCNBH3 , 95%), and
ferrocenecarboxaldehyde (FcCHO, 98%) were purchased from
Sigma–Aldrich Chemical (USA). All other chemicals were of analytical grade and used without further purification. All solutions
were prepared using doubly distilled water.
2.2. Instruments
X-ray diffraction (XRD) patterns of the nanocomposites were
carried out using a Rigaku powder diffractometer equipped with
Cu K␣1 radiation ( = 1.5406 Å). UV–vis absorption spectra were
recorded with a UV-2450 spectrophotometer (Shimadzu). Scanning
electron microscopy (SEM) images were obtained by using a Quanta
200 scanning electron microscope (FEI, USA). The accelerating voltage was 20 kV. The Fourier-transform infrared (FTIR) spectra of
samples in KBr pellets were recorded on a Nicolet 5700 FTIR spectrometer (Nicolet, USA). Electrochemical experiments were carried
out on a PGSTAT30/FRA2 system (Autolab, The Netherlands). A
three-electrode system including an Ag/AgCl (saturated KCl) reference electrode, a platinum wire as auxiliary electrode, and the
modified electrode as the working electrode was employed. The
electrolyte solutions were purged with N2 for at least 10 min to
remove O2 and kept under N2 atmosphere during measurements.
The electrochemical impedance spectroscopic (EIS) measurements
were performed at a bias potential of 210 mV in the presence of a
5.0 mM K3 [Fe(CN)6 ]/K4 [Fe(CN)6 ] (1:1) mixture as a redox probe in
0.1 M phosphate buffer (containing 0.1 M KCl, pH 6.98) by applying
an alternating current voltage with 5 mV amplitude in the frequency range from 0.01 Hz to 100 kHz.
2.3. Synthesis of graphene oxide
Graphene oxide (GO) was synthesized from graphite flake by a
modified Hummer’s method [46,47]. In brief, 0.5 g of graphite flake,
0.5 g of NaNO3 , and 23 mL of H2 SO4 were stirred together in an ice
bath. While maintaining vigorous agitation, an amount of 3 g of
KMnO4 was slowly added, and the rate of addition was controlled
Scheme 1. Schematic representation of the CS–Fc/GO/GOx modified electrode and the mechanism of the oxidation of glucose, catalyzed by GOx and mediated by Fc.
J.-D. Qiu et al. / Sensors and Actuators B 160 (2011) 287–294
289
Fig. 1. (A) XRD patterns of (a) pristine graphite and (b) GO. (B) SEM image and UV–vis spectrum of GO.
carefully to avoid a sudden increase in temperature. The mixture
was then maintained at 35 ◦ C for about 1 h. Deionized water (40 mL)
was gradually added, causing an increase in temperature to 90 ◦ C.
Finally, the mixture was further treated with 100 mL of deionized
water and 3 mL of 30% H2 O2 , turning the color of the solution from
dark brown to yellow. GO was then obtained through filtering,
water washing and drying process.
2.4. Preparation of ferrocene branched chitosan
CS–Fc was prepared according to the as-reported method [42].
Briefly, CS (75.0 mg) was dissolved in 0.1 M acetic acid solution
(15.0 mM). FcCHO (10.0 mg) was dissolved in methanol (10.0 mL)
and added to the CS solution. After the mixture was stirred at room
temperature for 2 h, NaCNBH3 (80.0 mg) was dropped and the reaction mixture was stirred for 24 h. The reaction was quenched by
precipitation with 5% NaOH and the yellow product was exhaustively washed with methanol and water. The product was dried in
air and finally redispersed in 0.2 M acetate buffer (pH 5.0) using
sonication.
2.5. Sensor construction
Glassy carbon electrode (GCE, diameter of 3 mm) was carefully
polished with 1.0, 0.3, and 0.05 ␮m alumina slurry, respectively,
and rinsed thoroughly with doubly distilled water between each
polishing step. The electrode was successively sonicated in 1:1
nitric acid, acetone, and doubly distilled water, and then allowed to
dry under N2 . Subsequently, 6 ␮L of the mixture consisting of CS–Fc
(0.5 mg mL−1 ), GO (0.03 mg mL−1 ), and GOx (4.0 mg mL−1 ) was cast
on the surface of well-polished GCE, and then dried in air at room
temperature. The obtained film was denoted as CS–Fc/GO/GOx. For
comparison, CS–Fc, CS–Fc/GO, and CS–Fc/GOx modified electrodes
were also prepared using the same procedure. All the resultant electrodes were stored at 4 ◦ C when not in use. The fabrication of the
CS–Fc/GO/GOx modified electrode and the redox-enzyme catalytic
cycle that gives rise to the current in the presence of glucose is
schematically shown in Scheme 1.
3. Results and discussion
3.1. Characterization of GO
The XRD patterns of the pristine graphite and GO were collected (Fig. 1A). The diffraction peak at 2␪ value of 26.5◦ (Fig. 1A,
curve a) in the XRD pattern of pristine graphite can be assigned
to the (0 0 2) facet of the hexagonal crystalline graphite [48–50].
Compared with the pristine graphite, disappearance of the peak at
about 26.5◦ and appearance of the peak at 10.0◦ (Fig. 1A, curve b)
reveal the successful oxidation of the starting graphite. SEM was
then used to investigate the surface morphological characters of
the as-synthesized GO (Fig. 1B). The isolated single-layer flakes
demonstrated that GO had been readily exfoliated into individual
sheets in water by ultrasonic treatment. To further demonstrate
the formation of GO, UV–vis spectroscopy was carried out (inset of
Fig. 1B). The characteristic absorption peak of GO at 231 nm confirmed that GO sheets were successfully synthesized, which was
consistent with other studies [31].
3.2. Characterization of the CS–Fc/GO/GOx biocomposite film
Fig. 2A shows a flat and featureless morphology of CS–Fc/GOx
film. When further entrapped GO into the CS matrix, as expected,
the formed CS–Fc/GO/GOx membrane displayed a rough surface
and many nanosheets uniformly dispersed throughout the hybrid
film without obvious aggregation (Fig. 2B), indicating that a biocomposite film composed of CS–Fc entrapped GO and GOx was
formed on the electrode surface. Fig. 2C and D represent the photographs of GO and CS–Fc/GO, respectively, in PBS buffer solutions
with pH ranging from 5.0 to 10.0. It can be observed from Fig. 2C that
the as-synthesized GO easily results in agglomeration in PBS solutions with different pHs, thus causing the color of solutions uneven
(Fig. 2C, a–e) [33]. In contrast, the CS–Fc/GO nanocomposite could
be readily dispersed in different PBS to form very homogenous dark
brown solutions (Fig. 2D, a–e) and no sediments were observed
for at least a month, which indicated that the nanocomposite was
highly hydrophilic. In addition, the CS–Fc functionalized GO could
be precipitated via high speed centrifuging and redispersed in fresh
water or buffer solution at least 6 times. These observations suggested that GO was efficiently prevented from aggregations in
PBS buffer solutions with different pHs by functionalizing with
CS–Fc.
3.3. FTIR spectroscopic analysis of the CS–Fc/GO/GOx composite
FTIR experiments were carried out to investigate the formation
of the CS–Fc/GO/GOx composite. Fig. 3 shows the FTIR spectra of
GO, CS–Fc, GO/CS–Fc, GOx, and GO/CS–Fc/GOx. In the spectrum
of GO (Fig. 3a), the peak at 1734 and 1622 cm−1 are characteristics of the C O stretch of the carboxylic group on the GO and
deformations of the O–H bond in water, respectively. In the spectrum of CS–Fc (Fig. 3b), there were two characteristic absorbance
bands centered at 1651 and 1596 cm−1 , which corresponded to
the C O stretching vibration of –NHCO– and the N–H bending
of –NH2 , respectively. Compared with pure CS–Fc and GO, both
peaks at 1596 cm−1 related to –NH2 absorbance vibration and
290
J.-D. Qiu et al. / Sensors and Actuators B 160 (2011) 287–294
Fig. 2. Typical SEM images of (A) CS–Fc/GOx and (B) CS–Fc/GO/GOx. Photographs of (C) pure GO and (D) CS–Fc/GO in PBS buffer solutions with different pHs (a) pH 5.0, (b)
pH 6.0, (c) pH 7.4, (d) pH 9.0, (e) pH 10.0.
at 1734 cm−1 belonging to C O stretch of the carboxylic group
disappeared in the spectra of CS–Fc/GO nanocomposite. Moreover,
the band corresponding to the C O characteristic stretching band
of the amide group shifted to a lower wavenumber (Fig. 3c). These
could be ascribed to the synergistic effect of hydrogen bonding
between CS–Fc and the oxygenated groups in GO and electrostatic interaction between polycationic CS and the negative charge
on the surface of GO [33]. After CS–Fc/GO further interacted with
GOx, it could be seen that the FTIR spectrum of the CS–Fc/GO/GOx
exhibited the amide I (1655 cm−1 ) and amide II (1585 cm−1 )
bands of the immobilized enzyme (Fig. 3d), which were essentially the same as those of the native GOx (Fig. 3e), demonstrated
the successful formation of CS–Fc/GO/GOx biocomposite and
the secondary structures of the immobilized enzyme were well
maintained.
a
b
The modified electrodes were monitored by EIS, which is an
effective method for probing the interfacial properties of the
electrodes [51] and often used for understanding chemical transformations and processes associated with the conductive supports
[52–54]. The typical impedance spectrum (presented in the form
of the Nyquist plot) includes a semicircle portion and a linear
portion. The semicircle portion at higher frequencies corresponds
to the electron-transfer limited process, and the linear portion
at lower frequencies represents the diffusion-limited process.
The semi-circle diameter in the impedance spectrum is equal to
the electron-transfer resistance, Ret , which controls the electrontransfer kinetics of the redox probe at the electrode interface. Fig. 4
shows the EIS results of a bare GCE, CS–Fc/GCE, CS–Fc/GO/GCE, and
1734 1622
1651
c
Transmittance(%)
3.4. Electrochemical impedance spectroscopic characterization of
the modified electrodes
1596
d
1640
e
1655
1585
1655
1585
4000
3500
3000
2500
2000
1500
-1
Wavenumber/ cm
Fig. 3. FTIR spectra of (a) GO, (b) CS–Fc, (c) GO/CS–Fc, (d) GO/CS–Fc/GOx, (e) GOx.
Fig. 4. The electrochemical impedance spectra of (a) bare GCE, (b) CS–Fc/GCE, (c)
CS–Fc/GO/GCE, and (d) CS–Fc/GO/GOx/GCE. Inset: equivalent circuit used to model
impedance data in the presence of the redox couples. Supporting electrolyte, 0.1 M
PBS (pH 6.98) +0.1 M KCl +5.0 mM Fe(CN)6 3−/4− solution.
J.-D. Qiu et al. / Sensors and Actuators B 160 (2011) 287–294
40
Current /μA
20
20
10
15
0
-10
-20
10
0
Current/μA
Current/μA
20
d
40
30
30
291
-30
-40
0
5
10 15 20 25 30
1/2
-1 1/2
V /(mVs )
-10
10
c
5
a
b
0
-20
-5
-30
-0.2 -0.1
0.0
0.1
0.2
0.3
0.4
0.5
0.6
-0.1 0.0
0.1
0.2
0.3
0.4
0.5
0.6
0.7
Potential/V
Potential/V
Fig. 5. CVs of CS–Fc/GO/GOx film electrode in 0.1 M PBS (pH 6.98, containing 0.1 M
KCl) at different scan rates (from inner to outer: 5, 10, 50, 100, 200, 300, 400, 500,
600,700, 800, 900, 1000 mV s−1 ). Inset: plots of peak current vs.1/2 .
CS–Fc/GO/GOx/GCE in the presence of redox probe Fe(CN)6 3−/4− . It
could be seen that the bare electrode exhibited an almost straight
line, which was the characteristic of a diffusion limiting step of the
electrochemical process (Fig. 4a). After coating a layer of CS–Fc, the
Ret of the modified electrode obviously increased to 95 (Fig. 4b),
which indicated the successful formation of CS–Fc film. When GO
was incorporated into CS–Fc film, the Ret of the resulted CS–Fc/GO
film surprisingly decreased to 26 (Fig. 4c), which was much
smaller than that of CS–Fc film, suggesting that the presence of
GO on the CS–Fc modified electrode formed high electron conduction pathways between the electrode and electrolyte, and greatly
improved the electron transfer of the redox probe. After immobilization of GOx on CS–Fc/GO film, Ret increased to 182 (Fig. 4d),
the increase of electron-transfer resistance confirmed that the GOx
was successfully assembled on the surface of GO, which resulted in
the hindered pathway of electron transfer. To obtain more detailed
information about the impedance of the modified electrode, a modified Randles equivalent circuit (inset in Fig. 4) was chosen to fit the
measured results. The two components of the scheme, Rs and Zw ,
represent the bulk properties of electrolyte solution and diffusion
of the applied redox probe, respectively. The other two components
of the circuit, Cdl and Ret , depend on the dielectric and insulating
features at the electrode/electrolyte interface.
3.5. Electrochemical behavior of the modified electrode
Fig. 5 displays typical cyclic voltammograms (CVs) of the
CS–Fc/GO/GOx film electrode in 0.1 M PBS (pH 6.98) over the potential range from −0.1 to 0.6 V at various scan rates. It is clear that a
couple of well-defined redox peaks of the immobilized Fc group
are observed at scan rates ranging from 5 to 1000 mV s−1 . Both the
reduction and oxidation peak currents increased obviously with
the increasing scan rates, while the peak-to-peak separation (E)
was nearly independent of the scan rates, which indicated that the
electron transfer of Fc in this composite film was relatively fast and
reversible. Meanwhile, the CS–Fc/GO/GOx modified electrode was
found to be extremely stable during continuous potential cycling
between −0.1 V and 0.6 V, indicated that the covalently bounded Fc
in CS chain could prevent mediator from leakage efficiently. Both
the anodic and cathodic peak currents increased linearly with the
square root of scan rate (1/2 ) (inset in Fig. 5), confirming that the
redox behavior of the electrodes was diffusion controlled, where
electron transferred to and from the redox centers of the CS–Fc
involved diffusion [55]. The linearity of ip against 1/2 plot also
Fig. 6. CVs of CS–Fc/GOx (a and c) and CS–Fc/GO/GOx (b and d) modified electrodes
at 50 mV s−1 in 0.1 M PBS (pH 6.98, containing 0.1 M KCl) in the absence (a and b)
and the presence (c and d) of 14 mM glucose.
suggested the electron-transfer model of Randles–Sevcik, which
assumes semi-infinite linear diffusion, the diffusion coefficient of
charge transfer (D) was determined in accordance to the equation
ip = (2.69 × 105 )n3/2 ACD1/2 1/2 [55,56]. Where ip is the peak current, n is the number of electrons, A is the electrode area (in cm2 ),
␯ is the scan rate (in V/s), D is the diffusion coefficient of charge
transfer, and C is the concentration of Fc redox centers in the film.
The product D was calculated 6.7 × 10−8 cm−2 s−1 , which is larger
than other redox polymers [57]. Electron-transfer under diffusion
control suggests that three-dimensional charge propagation in the
CS–Fc/GO/GOx film is operative with this redox polymer, unlike
the acrylamide–ferrocene copolymer reported previously in which
the redox charge decreased until a stable two-dimensional wired
enzyme was obtained in the monolayer level [58]. In the system
reported here, diffusion of the mediator and chain reptation are not
possible, thus charge propagate must through electron-hopping
between neighboring the conductive GO and polymer chain segments containing ferrocene redox couples.
3.6. Electrochemical response of CS–Fc/GO/GOx film to glucose
To investigate the role of GO in biosensors, CVs of the CS–Fc/GOx
and CS–Fc/GO/GOx modified electrodes in the absence and the
presence of glucose in 0.1 M PBS were studied (Fig. 6). In the absence
of glucose, the CS–Fc/GOx modified electrode exhibited a pair of
redox peaks of Fc (Fig. 6a). Further incorporation of GO in the film,
the anodic peak and cathodic peak currents of the CS–Fc/GO/GOx
modified electrode increased considerably (Fig. 6b). Upon addition
of glucose to the PBS solution, the anodic peak current increased
noticeably, while the cathodic peak current decreased, indicating
an obvious electrocatalytic oxidation of glucose at both CS–Fc/GOx
(Fig. 6c) and CS–Fc/GO/GOx (Fig. 6d) modified electrodes. However,
the electrocatalytic current on the CS–Fc/GO/GOx modified electrode was significantly improved, which was about 2.7 times higher
than that of the CS–Fc/GOx modified electrode. The improved performance of the CS–Fc/GO/GOx film could be attributed to the
presence of GO in the hybrid film. The nanoscale individual sheets
act as “molecular wires” to connect the active sites of GOx and
electron mediator Fc with the electrode, increasing the electron
transfer rate significantly. In addition, the high conductivity of GO
is also responsible for the increased current. This typical enzymedependent catalytic process shown in Scheme 1 can be further
expressed as follows [59]:
Glucose + GOx(FAD) → Gluconolactone + GOx(FADH2 )
(1)
292
J.-D. Qiu et al. / Sensors and Actuators B 160 (2011) 287–294
Table 1
Biosensor for glucose detection in comparison with literature.
app
Glucose Biosensor
Detection limit (␮M)
KM
(GOD/GNP/CS)6 /GOD/GNP/PAA/Pt
GOD/graphite/Nafion
pH polymer/BSA/catalase/GOx
GOD/MWNTs/CS–Fc
GOD/mesocellular carbon foam/Nafion
CS–Fc/GO/GOx
7.0
10
600
6.5
70
7.6
10.5
–
–
6.87
–
2.1
GOx(FADH2 ) + 2Fc+ → GOx(FAD) + 2Fc + 2H+
(2)
Fc ↔ Fc+ + e−
(3)
Where Fc and Fc+ represent reduced and oxidized forms of the ferrocene, GOx (FAD) and GOx (FADH2 ) are the oxidized and reduced
forms of glucose oxidase. In this process, the mediator takes the
place of oxygen in the native enzymatic reaction, and two electrons are transferred from glucose to the FAD of the enzyme. These
electrons can then be transferred from FADH2 to the mediator,
which is then oxidized at the electrode surface producing a current that is directly proportional to the concentration of glucose in
solution.
3.7. Amperometric determination of glucose
The current–time curve of the CS–Fc/GO/GOx film electrode
upon successive additions of glucose at an applied potential of
+0.3 V clearly illustrated that the modified electrode could respond
very rapidly to a change in the glucose concentration (Fig. 7), producing steady signals within only 5 s. Such a short response time
further proves that the CS–Fc/GO material is a promising platform
for the construction of biosensors. The response displayed a linear
glucose concentration range from 0.02 to 6.78 mM with a detection limit of 7.6 ␮M at a signal-to-noise ratio of 3. Based on three
times the standard deviation of the slope, CS–Fc/GO/GOx electrodes
exhibited an excellent sensitivity of 10 ␮A mM−1 cm−2 for glucose
detection, which was much higher than that of 7.86 ␮A mM−1 cm−2
for (GOD/GNP/CS)6 /GOD/GNP/PAA/Pt [60]. The Michaelis–Menten
app
constant (KM ), which is a reflection of both the enzymatic affinity and the ratio of microscopic kinetic constant, was calculated to
be 2.1 mM according to the Lineweaver–Burk equation [61]. The
further comparison of CS–Fc/GO/GOx film electrode developed in
this study with other biosensors based on GOx is shown in Table 1.
(mM)
Linearity (mM)
Reference
0.5–16
Upto 6
1–10
0.02–5.36
Upto 2
0.02–6.78
[60]
[62]
[63]
[64]
[65]
Current work
Results illustrate that the biosensor described in this work exhibited a lower detection limit and a wider linear range toward glucose
detection, thereby demonstrating the capability of the proposed
biosensor toward measuring the plasma glucose level (for the diagnosis of diabetes).
3.8. Reproducibility and stability of the biosensor
The reproducibility of the CS–Fc/GO/GOx electrode was evaluated using five enzyme electrodes prepared at identified conditions
on different days. A relative standard deviation (RSD) of the biosensor response to 1 mM glucose was 4.3%, indicating the satisfied
reproducibility of the resulting enzyme electrode for practical
application. The storage stability was investigated using the prepared electrode after having been stored in 0.1 M PBS (pH 6.98)
at 4 ◦ C for 1 month. No substantial decrease of the signal (>90%)
was obtained, which may be arising from the biocompatibility of
the CS–Fc/GO/GOx composite in maintaining the activity of the
enzyme.
3.9. Interference study
Fig. 8 presents the selectivity testing results of the
CS–Fc/GO/GOx modified electrode with successive additions
of ascorbic acid (AA), uric acid (UA), and glucose in 0.1 M pH 6.98
PBS solution. Interestingly, the as-prepared biosensor gave out
significant signals when adding 1 mM glucose for each time, while
there were essentially negligible current responses of 0.1 mM AA
and 0.02 mM UA at the normal physiological level. The results
obtained from several repetitions corroborated the fact that the
CS–Fc/GO/GOx modified electrodes would give higher sensitivity and selectivity for glucose detection under physiological
conditions.
6
5
i/μA
6
2.0
4
0
0
2
4
6
8
10
Cglucose/mM
3
0.5mM
2
0.04mM
1
Glucose
1.6
4
Current/μA
Current/μA
2
Glucose
Glucose
UA AA
Glucose
1.2
UAAA
0.8 Glucose
0.02mM
0.3mM
0
0.4
0.1mM
600
900
1200
1500
1800
2100
t/s
Fig. 7. Amperometric response of CS–Fc/GO/GOx modified electrode at applied
potential of +0.3 V to successive addition of glucose in 0.1 M PBS (pH 6.98, containing
0.1 M KCl).
700
800
900
1000
1100
1200
t/s
Fig. 8. Amperometric response of CS–Fc/GO/GOx modified electrode to successive
addition of 0.02 mM UA, 0.1 mM AA, and 1 mM glucose into stirred 0.1 M PBS (pH
6.98) containing 0.1 M KCl at a potential of +0.3 V.
J.-D. Qiu et al. / Sensors and Actuators B 160 (2011) 287–294
4. Conclusions
A novel ultrasensitive glucose biosensor by integrating
CS–Fc/GO nanocomposite as an ideal conductive platform for
the enzyme immobilization was designed through a simple selfassembly method. The uniform GO dispersion within the CS matrix
could significantly improve the stability of GO and make GO exhibit
a positive charge, which was more favorable for the further immobilization of negative charged GOx without destructing its native
structure and bioactivity. Further attaching redox mediator ferrocene group to CS matrix could not only effectively prevent the
leakage of Fc from the matrix and improve the electrical conductivity of CS, but also show excellent biocompatibility for enzyme
immobilization. The proposed CS–Fc/GO/GOx electrode performed
good electrocatalytic oxidation for glucose with a broad linearity, good sensitivity, excellent reproducibility and storage stability.
Therefore, such a multicomponent platform, which integrates the
advantages of GO, ferrocene, chitosan, and GOx together, may have
promising potential for the fabrication of high performance flexible
biosensors or bioreactors.
Acknowledgments
This work was supported by grants from the National Natural
Science Foundation of China (20865003, 20805023, 21065006) and
the Natural Science Foundation of Jiangxi Province (2010GZH0094).
References
[1] B.R. Azamian, J.J. Davis, K.S. Coleman, C.B. Bagshaw, M.L.H. Green, Bioelectrochemical single-walled carbon nanotubes, J. Am. Chem. Soc. 124 (2002)
12664–12665.
[2] Y.H. Lin, W. Yantasee, J. Wang, Carbon nanotubes (CNTs) for the development
of electrochemical biosensors, Front. Biosci. 10 (2005) 492–505.
[3] G.A. Rivas, M.D. Rubianes, M.C. Rodriguez, N.E. Ferreyra, G.L. Luque, M.L. Pedano,
S.A. Miscoria, C. Parrado, Carbon nanotubes for electrochemical biosensing,
Talanta 74 (2007) 291–307.
[4] P. Yanez-Sedeno, J.M. Pingarron, Gold nanoparticle-based electrochemical
biosensors, Anal. Bioanal. Chem. 382 (2005) 884–886.
[5] B.K. Jena, C.R. Raj, Electrochemical biosensor based on integrated assembly
of dehydrogenase enzymes and gold nanoparticles, Anal. Chem. 78 (2006)
6332–6339.
[6] J.M. Pingarron, P. Yanez-Sedeno, A. Gonzalez-Cortes, Gold nanoparticle-based
electrochemical biosensors, Electrochim. Acta 53 (2008) 5848–5866.
[7] Y. Yang, X. Yang, C.X. Jiao, H.F. Yang, Z.M. Liu, G.L. Shen, R.Q. Yu, Optical sensor for berberine utilizing its intrinsic fluorescence enhanced by the formation
of inclusion complex with butylated-beta-cyclodextrin, Anal. Chim. Acta 513
(2004) 385–392.
[8] I.M. Hsing, Y. Xu, W.T. Zhao, Micro- and nano-magnetic particles for applications in biosensing, Electroanalysis 19 (2007) 755–768.
[9] W. Vastarella, R. Nicastri, Enzyme/semiconductor nanoclusters combined systems for novel amperometric biosensors, Talanta 66 (2005) 627–633.
[10] A. Heller, B. Feldman, Electrochemical glucose sensors and their applications
in diabetes management, Chem. Rev. 108 (2008) 2482–2505.
[11] Y.J. Zou, C.L. Xiang, L.X. Sun, F. Xu, Glucose biosensor based on electrodeposition
of platinum nanoparticles onto carbon nanotubes and immobilizing enzyme
with chitosan–SiO2 sol–gel, Biosens. Bioelectron. 23 (2008) 1010–1016.
[12] M.D. Shirsat, C.O. Too, G.G. Wallace, Amperometric glucose biosensor on layer
by layer assembled carbon nanotube and polypyrrole multilayer film, Electroanalysis 20 (2008) 150–156.
[13] V. Vamvakaki, K. Tsagaraki, N. Chaniotakis, Carbon nanofiber-based glucose
biosensor, Anal. Chem. 78 (2006) 5538–5542.
[14] Y. Du, X.L. Luo, J.J. Xu, H.Y. Chen, A simple method to fabricate a chitosan–gold
nanoparticles film and its application in glucose biosensor, Bioelectrochemistry
70 (2007) 342–347.
[15] L. Li, Q.L. Sheng, J.B. Zheng, H.F. Zhang, Facile and controllable preparation of
glucose biosensor based on prussian blue nanoparticles hybrid composites,
Bioelectrochemistry 74 (2008) 170–175.
[16] X.W. Sun, J.X. Wang, A. Wei, Zinc oxide nanostructured biosensor for glucose
detection, J. Mater. Sci. Technol. 24 (2008) 649–656.
[17] A. Kaushika, R. Khan, P.R. Solanki, P. Pandey, J. Alam, S. Ahmad, B.D. Malhotra,
Iron oxide nanoparticles-chitosan composite based glucose biosensor, Biosens.
Bioelectron. 24 (2008) 676–683.
[18] J.J. Yu, D.L. Yu, T. Zhao, B.Z. Zeng, Development of amperometric glucose biosensor through immobilizing enzyme in a Pt nanoparticles/mesoporous carbon
matrix, Talanta 74 (2008) 1586–1591.
293
[19] M. Pumera, Electrochemistry of graphene: new horizons for sensing and energy
storage, Chem. Rec. 9 (2009) 211–223.
[20] M. Pumera, Graphene-based nanomaterials and their electrochemistry, Chem.
Soc. Rev. 39 (2010) 4146–4157.
[21] S. Alwarappan, A. Erdem, C. Liu, C.Z. Li, Probing the electrochemical properties
of graphene nanosheets for biosensing applications, J. Phys. Chem. C 113 (2009)
8853–8857.
[22] G. Li, H. Xu, W.J. Huang, Y. Wang, Y.S. Wu, R. Parajuli, A pyrrole quinoline
quinone glucose dehydrogenase biosensor based on screen-printed carbon
paste electrodes modified by carbon nanotubes, Meas. Sci. Technol. 19 (2008)
065203.
[23] A. Kusakari, M. Izumi, H. Ohnuki, Preparation of an enzymatic glucose sensor based on hybrid organic–inorganic Langmuir–Blodgett films: adsorption of
glucose oxidase into positively charged molecular layers, Colloids Surf. Physicochem. Eng. Aspects 321 (2008) 47–51.
[24] Y. Wang, Y.M. Li, L.H. Tang, J. Lu, J.H. Li, Application of graphene-modified electrode for selective detection of dopamine, Electrochem. Commun. 11 (2009)
889–892.
[25] M. Pumera, A. Ambrosi, A. Bonanni, E.L.K. Chng, H.L. Poh, Graphene for electrochemical sensing and biosensing, Trends Anal. Chem. 29 (2010) 954–965.
[26] T.H. Seah, M. Pumera, Platelet graphite nanofibers/soft polymer composites for
electrochemical sensing and biosensing, Sens. Actuators B: Chem. 156 (2011)
79–83.
[27] S. Stankovich, D.A. Dikin, R.D. Piner, K.A. Kohlhaas, A. Kleinhammes, Y. Jia, Y. Wu,
S.T. Nguyen, R.S. Ruoff, Synthesis of graphene-based nanosheets via chemical
reduction of exfoliated graphite oxide, Carbon 45 (2007) 1558–1565.
[28] M.J. McAllister, J.L. Li, D.H. Adamson, H.C. Schniepp, A.A. Abdala, J. Liu, M.
Herrera-Alonso, D.L. Milius, R. Car, R.K. Prud’homme, I.A. Aksay, Single sheet
functionalized graphene by oxidation and thermal expansion of graphite,
Chem. Mater. 19 (2007) 4396–4404.
[29] S. Stankovich, R.D. Piner, X.Q. Chen, N.Q. Wu, S.T. Nguyen, R.S. Ruoff, Stable
aqueous dispersions of graphitic nanoplatelets via the reduction of exfoliated
graphite oxide in the presence of poly(sodium 4-styrenesulfonate), J. Mater.
Chem. 16 (2006) 155–158.
[30] S. Niyogi, E. Bekyarova, M.E. Itkis, J.L. McWilliams, M.A. Hamon, R.C. Haddon,
Solution properties of graphite and graphene, J. Am. Chem. Soc. 128 (2006)
7720–7721.
[31] D. Li, M.B. Muller, S. Gilje, R.B. Kaner, G.G. Wallace, Processable aqueous dispersions of graphene nanosheets, Nat. Nanotechnol. 3 (2008) 101–105.
[32] Q.A. Zhang, Y. Qiao, F. Hao, L. Zhang, S.Y. Wu, Y. Li, J.H. Li, X.M. Song, Fabrication of a biocompatible and conductive platform based on a single-stranded
DNA/graphene nanocomposite for direct electrochemistry and electrocatalysis,
Chem. Eur. J. 16 (2010) 8133–8139.
[33] A. Saha, S.K. Basiruddin, S.C. Ray, S.S. Roy, N.R. Jana, Functionalized graphene
and graphene oxide solution via polyacrylate coating, Nanoscale 2 (2010)
2777–2782.
[34] X. Yang, Y. Tu, L. Li, S. Shang, X.M. Tao, Well-dispersed chitosan/graphene oxide
nanocomposites, ACS Appl. Mater. Interfaces 2 (2010) 1707–1713.
[35] A.E.G. Cass, G. Davis, G.D. Francis, H.A.O. Hill, W.J. Aston, I.J. Higgins, E.V. Plotkin,
L.D.L. Scott, A.P.F. Turner, Ferrocene-mediated enzyme electrode for amperometric determination of glucose, Anal. Chem. 56 (1984) 667–671.
[36] J. Razumiene, A. Vilkanauskyte, V. Gureviciene, V. Laurinavicius, N.V. Roznyatovskaya, Y.V. Ageeva, M.D. Reshetova, A.D. Ryabov, New bioorganometallic
ferrocene derivatives as efficient mediators for glucose and ethanol biosensors
based on PQQ-dependent dehydrogenases, J. Organomet. Chem. 668 (2003)
83–90.
[37] M. Zayats, E. Katz, R. Baron, I. Willner, Reconstitution of apo-glucose dehydrogenase on pyrroloquinoline quinone-functionalized Au nanoparticles yields
an electrically contacted biocatalyst, J. Am. Chem. Soc. 127 (2005) 12400–
12406.
[38] E. Mikeladze, A. Schulte, M. Mosbach, A. Blochl, E. Csoregi, R. Solomonia, W.
Schumann, Redox hydrogel-based bienzyme microelectrodes for amperometric monitoring of l-glutamate, Electroanalysis 14 (2002) 393–399.
[39] A. Belay, A. Collins, T. Ruzgas, P.T. Kissinger, L. Gorton, E. Csoregi, Redox hydrogel
based bienzyme electrode for l-glutamate monitoring, J. Pharm. Biomed. Anal.
19 (1999) 93–105.
[40] S. Cosnier, A. Novoa, C. Mousty, R.S. Marks, Biotinylated alginate immobilization
matrix in the construction of an amperometric biosensor: application for the
determination of glucose, Anal. Chim. Acta 453 (2002) 71–79.
[41] S. Jawaheer, S.F. White, S. Rughooputh, D.C. Cullen, Enzyme stabilization using
pectin as a novel entrapment matrix in biosensors, Anal. Lett. 35 (2002)
2077–2091.
[42] J.D. Qiu, R. Wang, R.P. Liang, X.H. Xia, Electrochemically deposited nanocomposite film of CS–Fc/Au NPs/GOx for glucose biosensor application, Biosens.
Bioelectron. 24 (2009) 2920–2925.
[43] W.W. Yang, H. Zhou, C.Q. Sun, Synthesis of ferrocene-branched chitosain derivatives: redox polysaccharides and their application to reagentless
enzyme-based biosensors, Macromol. Rapid Commun. 28 (2007) 265–270.
[44] A. Garcia, C. Peniche-Covas, B. Chico, B.K. Simpson, R. Villalonga, Ferrocene
branched chitosan for the construction of a reagentless amperometric hydrogen peroxide biosensor, Macromol. Biosci. 7 (2007) 435–439.
[45] S. Wu, Y. Chen, F. Zeng, S. Gong, Z. Tong, Electron transfer in ferrocenecontaining functionalized chitosan and its electrocatalytic decomposition of
peroxide, Macromolecules 39 (2006) 6796–6799.
[46] W.S. Hummers, R.E. Offeman, Preparation of graphitic oxide, J. Am. Chem. Soc.
80 (1958) 1339.
294
J.-D. Qiu et al. / Sensors and Actuators B 160 (2011) 287–294
[47] L.J. Cote, F. Kim, J. Huang, Langmuir–Blodgett assembly of graphite oxide single
layers, J. Am. Chem. Soc. 131 (2008) 1043–1049.
[48] Y.G. Zhou, J.J. Chen, F.B. Wang, Z.H. Sheng, X.H. Xia, A facile approach to the synthesis of highly electroactive Pt nanoparticles on graphene as an anode catalyst
for direct methanol fuel cells, Chem. Commun. 46 (2010) 5951–5953.
[49] H.L. Guo, X.F. Wang, Q.Y. Qian, F.B. Wang, X.H. Xia, A green approach to the
synthesis of graphene nanosheets, ACS Nano 3 (2009) 2653–2659.
[50] H.K. Jeong, Y.P. Lee, R.J.W.E. Lahaye, M.H. Park, K.H. An, I.J. Kim, C.W. Yang, C.Y.
Park, R.S. Ruoff, Y.H. Lee, Evidence of graphitic AB stacking order of graphite
oxides, J. Am. Chem. Soc. 130 (2008) 1362–1366.
[51] R. Ehret, W. Baumann, M. Brischwein, A. Schwinde, K. Stegbauer, B. Wolf, Monitoring of cellular behaviour by impedance measurements on interdigitated
electrode structures, Biosens. Bioelectron. 12 (1997) 29–41.
[52] S.Y. Dong, G. Luo, J. Feng, Q.W. Li, H. Gao, Immunoassay of staphylococcal
enterotoxin C1 by FTIR spectroscopy and electrochemical gold electrode, Electroanalysis 13 (2001) 30–33.
[53] A. Bardea, E. Katz, I. Willner, Probing antigen–antibody interactions on electrode supports by the biocatalyzed precipitation of an insoluble product,
Electroanalysis 12 (2000) 1097–1106.
[54] C.L. Feng, Y.H. Xu, L.M. Song, Study on highly sensitive potentiometric IgG
immunosensor, Sens. Actuators B: Chem. 66 (2000) 190–192.
[55] D. Shan, W. Yao, H. Xue, Electrochemical study of ferrocenemethanol-modified
layered double hydroxides composite matrix: application to glucose amperometric biosensor, Biosens. Bioelectron. 23 (2007) 432–437.
[56] E.J. Calvo, R. Etchenique, C. Danilowicz, L. Diaz, Electrical communication
between electrodes and enzymes mediated by redox hydrogels, Anal. Chem.
68 (1996) 4186–4193.
[57] J. Chen, K. Aoki, T. Nishiumi, T. Li, Voltammetry of suspensions of hollow particles with ferrocene-immobilized polyallylamine shells, Langmuir 22 (2006)
10510–10514.
[58] E.J. Calvo, C. Danilowicz, L. Diaz, Enzyme catalysis at hydrogel-modified electrodes with redox polymer mediator, J. Chem. Soc.: Faraday Trans. 89 (1993)
377–384.
[59] S.A. Merchant, M.T. Meredith, T.O. Tran, D.B. Brunski, M.B. Johnson, D.T.
Glatzhofer, D.W. Schmidtke, Effect of mediator spacing on electrochemical and
enzymatic response of ferrocene redox polymers, J. Phys. Chem. C 114 (2010)
11627–11634.
[60] B.Y. Wu, S.H. Hou, F. Yin, J. Li, Z.X. Zhao, J.D. Huang, Q. Chen, Amperometric glucose biosensor based on layer-by-layer assembly of multilayer films composed
of chitosan, gold nanoparticles and glucose oxidase modified Pt electrode,
Biosens. Bioelectron. 22 (2007) 838–844.
[61] R.A. Kamin, G.S. Wilson, Rotating ring-disk enzyme electrode for biocatalysis
kinetic studies and characterization of the immobilized enzyme layer, Anal.
Chem. 52 (1980) 1198–1205.
[62] J. Lu, L.T. Drzal, R.M. Worden, I. Lee, Simple fabrication of a highly sensitive glucose biosensor using enzymes immobilized in exfoliated graphite nanoplatelets
nafion membrane, Chem. Mater. 19 (2007) 6240–6246.
[63] Q. Cai, K. Zeng, C. Ruan, T.A. Desai, C.A. Grimes, A wireless, remote query glucose
biosensor based on a pH-sensitive polymer, Anal. Chem. 76 (2004) 4038–4043.
[64] R.P. Liang, L.X. Fan, R. Wang, J.D. Qiu, One-step electrochemically deposited
nanocomposite film of CS–Fc/MWNTs/GOD for glucose biosensor application,
Electroanalysis 21 (2009) 1685–1691.
[65] D. Lee, J. Lee, J. Kim, H.B. Na, B. Kim, C.H. Shin, J.H. Kwak, A. Dohnalkova, J.W.
Grate, T. Hyeon, H.S. Kim, Simple fabrication of a highly sensitive and fast glucose biosensor using enzymes immobilized in mesocellular carbon foam, Adv.
Mater. 17 (2005) 2828–2833.
Biographies
Jian-Ding Qiu is a professor in Department of Chemistry, Nanchang University, PR
China. He received his PhD in analytical chemistry from Zhongshan University of
China in 2004. He served as a postdoctoral research associate at the Nanjing University (2004–2006) and Hokkaido University (2009–2010), respectively. His current
research interests include bioelectroanalysis, nanotechnology, and bioinformatics.
Jing Huang is a MS candidate in Department of Chemistry, Nanchang University.
Her current research is the fabrication of a biosensor and its application in biological
fields.
Ru-Ping Liang received her PhD degree from Zhongshan University in 2004.
Presently, she is a professor in Department of Chemistry, Nanchang University.
Her main research interests include chemical modified electrode, microfluidics, and
surface modification.