Accepted Manuscript Detection of glucose using immobilised bi-enzyme on cyclic bisureas-gold nanoparticle conjugate Manjusha Mathew, N. Sandhyarani PII: DOI: Reference:
S0003-2697(14)00198-5 http://dx.doi.org/10.1016/j.ab.2014.05.003 YABIO 11736
To appear in:
Analytical Biochemistry
Received Date: Revised Date: Accepted Date:
10 February 2014 29 April 2014 6 May 2014
Please cite this article as: M. Mathew, N. Sandhyarani, Detection of glucose using immobilised bi-enzyme on cyclic bisureas-gold nanoparticle conjugate, Analytical Biochemistry (2014), doi: http://dx.doi.org/10.1016/j.ab. 2014.05.003
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1
Detection of glucose using immobilised bi-enzyme on cyclic
2
bisureas-gold nanoparticle conjugate
3 4 5 6 7 8
Manjusha Mathew, N. Sandhyarani* Nanoscience Research Laboratory, School of Nano Science and Technology, National Institute of Technology Calicut, Calicut, Kerala, India *Author to whom correspondence to be addressed. e-mail: [email protected] Ph: 91-495 2286537, Fax: 91-495 2287250
9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24
1
25
Abstract
26 27
A highly sensitive electrochemical glucose sensor has been developed by the co-
28
immobilization of glucose oxidase (GOx) and horseradish peroxidase (HRP) onto a gold
29
electrode modified with biocompatible cyclic bisureas-gold nanoparticle conjugate (CBU-
30
AuNP). A self-assembled monolayer of mercapto propionic acid (MPA) and CBU-AuNP was
31
formed on the gold electrode through a layer by layer assembly. This modified electrode was
32
used for immobilisation of the enzymes GOx and HRP. Both the HRP and GOx retained their
33
catalytic activity for an extended time indicated by the low value of Michaelis-Menten
34
constant. Analytical performance of the sensor was examined in terms of sensitivity,
35
selectivity, reproducibility, lower detection limit and stability. The developed sensor surface
36
exhibited a limit of detection of 100nM with a linear range of 100nM to 1mM. A high
37
sensitivity of 217.5µAmM-1cm-2 at a low potential of -0.3V was obtained in this sensor
38
design. Various kinetic parameters were calculated. The sensor was examined for its practical
39
clinical application by estimating glucose in human blood sample.
40 41 42
Keywords- glucose sensor; enzyme; glucose oxidase; horseradish peroxidase; cyclic
43
voltammetry
44 45 46 47 48 49
2
50
1. Introduction
51 52
Accurate determination and regular monitoring of blood glucose concentration is very
53
important in the diagnosis and treatment of diabetes mellitus. In addition, sensitive and
54
accurate detection of glucose has extensive applications in food industry and biotechnology
55
[1,2]. Specific recognition of glucose by glucose oxidase and reactions of glucose by glucose
56
oxidase are widely explored in commonly used electrochemical glucose sensors [3-5]. Since
57
the flavine adenine dinucleotide (FAD) redox centre of GOx is deeply buried inside a
58
protective protein shell, it is difficult to establish a direct electron communication between
59
the enzyme and electrode. In order to improve the sensitivity, mediated electron transfer
60
using ferrocene derivatives [6,7] and electroactive polymers [8] were reported. Application of
61
mediators provides higher sensitivity, however their application is limited as they
62
contaminate the sample solution. Consequently, modification of electrode surface to allow
63
efficient electron transfer between the active site of redox protein and electrode become
64
significant in the development of biosensors. Formation of well-ordered, homogeneous and
65
molecular monolayers on the surface of electrode by self assembly [9] can facilitate electron
66
transfer between the analyte and the electrode, offering good reproducibility with a fast
67
response [10]. In addition, integration of nanostructured materials with enzymatic biosensor
68
is widely used to improve the electron transfer between the GOx and electrode [11-13].
69
Enzymes are highly sensitive and may undergo denaturation with variation in conditions like
70
temperature, pH, humidity etc. Hence, enzyme attachment to the solid substrate and its
71
compatibility with the metal electrode is crucial in the realization of a sensitive biosensor.
72
Consequently, a biocompatible environment is needed to sustain the enzyme activity for an
73
extended time.
3
74
Co-immobilization of HRP with a H2O2 producing oxidase enzyme to fabricate a bi-
75
enzyme biosensor is an effective way to improve the glucose sensor performance [14-19]. In
76
this system, the H2O2 produced as a result of the enzyme induced oxidation of glucose is
77
subsequently reduced by HRP which permits the detection of glucose at a low potential. Thus
78
the bi-enzyme system switches the detection principle from an electrochemical oxidation
79
reaction to reduction reaction.
80
We have reported the efficient use of CBU-AuNP for the development of a hydrogen
81
peroxide sensor. Owing to the biocompatible environment of CBU and electron transfer
82
capability of gold nanoparticles we achieved one of the lowest reported limits of detection for
83
hydrogen peroxide [20]. The biocompatible environment and efficient electron transfer
84
ability provided by the CBU-AuNP have also been employed in the development of a glucose
85
sensor. On using the immobilized GOx on the surface we could detect the glucose only up to
86
1 M which is in the similar range as that of most of the GOx based glucose sensors [21-23].
87
To improve the detection limit of glucose sensor we have used a bi-enzyme strategy and
88
achieved a nanomolar level detection limit for glucose.The work reported in this manuscript
89
demonstrates the development of a highly sensitive glucose sensor using the immobilization
90
of GOx and HRP- on a gold electrode modified with CBU capped gold nanoparticles. H2O2
91
produced as a result of glucose oxidation by GOx served as a substrate for the HRP.
92
Optimized conditions of sensor performance were studied. Selectivity, reproducibility and
93
stability of the sensor surface were also evaluated and the sensor was used for the
94
measurement of the glucose level in blood.
95
2. Experimental Methods
96
2.1. Materials
97 98
GOx from Aspergillus niger, Peroxidase from horseradish type V1-A, hydrogen tetrachloroaurate trihydrate, 1,6-diisocyanato hexane and L-cystine
4
dimethyl ester
99
dihydrochloride were
obtained from Sigma-Aldrich. Monosodium glutamate, dextrose
100
anhydrous, Na2CO3, Na2HPO4 and KH2PO4 were purchased from Merck international. The
101
supporting electrolyte 0.1M Phosphate Buffered Saline buffer was prepared by mixing the
102
stock solutions of 0.1M Na2HPO4 and KH2PO4 and the pH was adjusted using NaOH or HCl.
103
All solutions were prepared with ultrapure water obtained from an ultrafiltration system
104
(18M cm. Milli-Q, Millipore system).
105 106
2.2. Synthesis
107 108
A standard procedure was adopted for the synthesis of CBU [24]. The hydrogen
109
bonding possibility of CBU has been employed for the immobilization of enzymes. To
110
achieve efficient electron transport between the gold electrode and analyte the CBU was
111
conjugated to gold nanoparticles. CBU-AuNP was prepared as described elsewhere [20].
112 113
2.3. Preparation of electrodes
114 115
Gold electrode of geometric area 0.28cm2 was mechanically polished with 0.3μm -
116
Al2O3 slurry and washed with distilled water. Then the electrode was sonicated in a freshly
117
prepared Piranha solution for 10min. It was then washed with deionised water, sonicated and
118
rinsed with ethanol and dried. A monolayer of MPA was prepared by immersing the cleaned
119
gold electrode in 1mM MPA in ethanol for 12h at room temperature. The electrode was taken
120
out and thoroughly rinsed with ethanol to remove the physically adsorbed thiol and dried.
121
Then the electrode was placed into CBU-AuNP composite in methanol for 4h. Finally
122
modification by the enzymes to the electrode surface was attained by incubating the electrode
5
123
in a mixture of 0.1mM GOx and 0.1mM HRP in PBS for 12h. After thoroughly rinsing the
124
sensor with deionised water, it was kept at 4°C when not in use.
125 126
2.4. Characterizations
127 128
Cyclic voltammetry (CV), chronoamperometry and quartz crystal microbalance -
129
QCM (Resonance frequency of the crystal - 7.9MHz) measurements were performed using
130
CH 400A Electrochemical Quartz Crystal Microbalance (CH Instruments, Austin, Texas). A
131
three electrode system was used with the modified gold electrode as working electrode, a
132
platinum wire as auxiliary electrode and saturated calomel as the reference electrode against
133
which all potentials were measured. The measurements have been carried out at room
134
temperature with a scan rate of 100mV/s. QCM experiments were carried out on a gold
135
polished quartz crystal electrode as a working electrode, calomel electrode as reference
136
electrode and a platinum wire as counter electrode. Atomic force microscopy (AFM)
137
experiments were performed using Park XE-100 atomic force microscope. Typical AFM
138
images were acquired in non contact mode with silicon nitride cantilever. Energy dispersive
139
X-ray spectroscopy (EDS) measurements were performed using
140
pressure field emission scanning electron microscope at an accelerating voltage 15kV.
141
Contact angle measurements were performed with a stereo zoom microscope. A micropipette
142
was used to dispense 10μL of millipore water on electrode surface and placed under the
143
objective. All measurements were carried out at room temperature.
144 145 146 147
6
Hitachi SU6600 variable
148
3.
Results and discussion
149 150
A schematic representation of the sensor surface is shown in figure 1. Initially the
151
MPA monolayers were formed on the cleaned gold electrode which was then used for the
152
assembly of CBU-AuNP. On this layer, GOx and HRP have been immobilised by taking
153
advantage of the large hydrogen bonding possibility of CBU.
154 155
Fig.1. Schematic representation of Au/MPA/CBU-AuNP/GOx-HRP sensor surface on each
156
stages of fabrication (A) MPA monolayer on Au (B)Assembled CBU-AuNP on MPA
157
monolayer (C)Immobilised enzymes on the CBU-AuNP surface.
158 159
3.1. Characterization of electrode surface
160 161
3.1.1. Cyclic voltammetry
162 163
Formation of each layer on the electrode surface was monitored using the cyclic
164
voltammogram of ferro/ferricyanide redox couple. The experiment was conducted in the
165
electrolyte containing 5mM K3[Fe(CN)6] and 0.1M KCl at a scan rate of 100mVs-1 for four
166
different types of electrodes viz. bare Au, Au/MPA, Au/MPA/CBU-AuNP and
167
Au/MPA/CBU-AuNP/GOx-HRP. Figure 2 compares the voltammetric responses at each
168
stages of the fabrication process.
169 170
Fig.2. Cyclic voltammograms of
bare Au,
171
Au/MPA/CBU-AuNP/GOx-HRP electrodes in 0.1M KCl solution containing 5mM
172
K3[Fe(CN)6].
7
Au/MPA, Au/MPA/CBU-AuNP and
173 174
A well defined redox peak of the Fe2+/Fe3+ was observed on bare gold electrode.
175
Modification of the electrode surface with MPA causes a decrease in the anodic and cathodic
176
current as it forms an insulating layer on the electrode, which acted as a blockade to the
177
electron transfer between the analyte and electrode surface. A remarkable increase of current
178
is obtained after CBU-AuNP assembly due to the electron transfer through the gold
179
nanoparticles. In addition to the electron transfer property, the high surface area of gold
180
nanoparticles and its enhanced catalytic activity [25,26] also plays an important role in
181
enhancing the current. Immobilization of enzymes on the electrode decreases the redox
182
current values due to the insulating nature of the enzymes.
183 184
3.1.2. Electrochemical quartz crystal microbalance
185 186
The surface modification of electrode by various immobilization steps are also confirmed
187
by monitoring the mass change in electrochemical quartz crystal microbalance (EQCM). The
188
change in resonance frequency of the crystal due to the increase in mass as a result of
189
adsorption of different materials on the electrode surface was recorded. The relationship
190
between the changes in mass per unit area (Δm) and frequency (Δf) are given by Sauerbrey
191
equation [27];
192 193
(1) where,
,
,
,
,
and
represent the frequency change, resonance frequency of
194
the fundamental mode(Hz), mass change(ng), area of the crystal (0.205cm2), shear
195
modulus(g/cms2) and density(g/cm3) of the crystal respectively.
196
EQCM results indicate that formation of different layers resulted in a decrease in
197
frequency and an increase in mass on the surface. The immobilization of MPA induces a shift 8
198
in the frequency of 25Hz, which corresponds to the adsorbed mass of 34.75 ng. The surface
199
concentration of MPA is calculated as 9.6×1014 molecules/cm2. Immobilization of CBU-
200
AuNP leads to a deposition of 201.55ng on the quartz crystal. Deposition of enzymes resulted
201
in a frequency change of 89Hz corresponding to a deposition of 123.71ng of enzymes.
202
(Molecular weight of GOx is 160kD and HRP is 44kD). The surface concentration of enzyme
203
is obtained as 1.7×1012 molecules/cm2 assuming equal probability of immobilization of both
204
the enzymes leading to equal concentration of enzymes on the surface. Table 1 summarizes
205
the quartz crystal frequency change and the mass adsorbed on the crystal surface after each
206
steps of immobilization.
207 208
Table.1. Quartz crystal frequency change and the mass adsorbed on the crystal surface after
209
the immobiliztion of MPA, CBU-AuNP, GOx-HRP
210 211
EDS is an effective tool to probe the change in weight percentage of elements on each
212
stages of immobilization. Table S1 of supplementary information summarizes the weight
213
percentage obtained on successive immobilization of different layers on the electrode surface.
214
Increases in weight percentages of carbon, sulphur and nitrogen confirm the successful
215
immobilization of each layer on the electrode surface.
216
Contact angle of the sensor surface on each stages of immobilization was also
217
measured. Microscopic images of a water droplet placed on each layer of Au/MPA/CBU-
218
AuNP/GOx-HRP electrode are given in figure S1 of supplementary information. A bare gold
219
electrode showed a contact angle of 72.8°. A decrease in contact angle to 62° is observed on
220
immobilization of MPA indicating an increased wetting due to the terminating carboxyl
221
group. Presence of biocompatible gold nanoparticles further increased the hydrophilicity of
222
the surface leading to a decrease in the contact angle value to 56°. Immobilization of enzymes
9
223
on the electrode leads to an additional reduction in the contact angle to 43.6°. The increase in
224
hydrophilic nature confirms the presence of enzymes on the electrode surface. The increased
225
hydrophilicity of gold nanoparticles helps to retain the activity of enzyme on the electrode
226
surface. Atomic force microscopic images also were taken to confirm the immobilization of
227
each step which is presented in supplementary information (figure S2).
228 229
3.2. Electrocatalytic oxidation of glucose at the bi-enzyme electrode
230 231
Optimum response of the electrode surface under different concentrations of the enzymes
232
were investigated and found that the molar ratio used for immobilization was critical in
233
determining the experimentally observed lower detection limit (ldl) [28]. Optimum
234
concentrations of [GOx]/[HRP] were determined by measuring the sensor performance by
235
varying the molar ratios of enzymes used for immobilisation. The current response and lower
236
detection limit in each case were determined. Three different molar ratios of GOx and HRP
237
viz. 1:1, 2:1 and 1:2 were selected for the preliminary experiments and the observed ldl was
238
of 100nM in the case of 1:1 and 1:2 and 1μM in the case of 2:1 ratio of GOx and HRP. Figure
239
S3A shows a comparison of the limit of detection of the three sensor surfaces. Since the 1:1
240
and 1:2 ratios of GOx and HRP exhibited the same ldl, we varied the concentration of HRP
241
from 1 to 8 times keeping the GOx concentration constant and determined the ldl. The lowest
242
detection limit was achieved with molar ratios of 1:1 and 1:2 of GOx and HRP respectively.
243
The result is given in the supplementary information as figure S3B. The current response
244
obtained for the oxidation of 1mM glucose with different molar ratios of [GOx]/[HRP] were
245
monitored by the change in current of H2O2 reduction. Current values of 90µA, 50µA, 49µA
246
and 35µA were obtained for the GOx:HRP concentration ratios of 1:1, 1:2, 1:4 and 1:8
247
respectively. Since the electrode prepared with GOx and HRP at a molar ratio of 1:1 yielded
10
248
maximum current response for 1mM glucose, we selected 1:1 molar ratio of GOx and HRP as
249
the optimum concentration and the response of the resulted electrode was studied in detail.
250
Figure 3A shows the response of the sensor at various concentrations of glucose. An
251
increase in cathodic current was observed with increase in concentration of glucose.
252
Minimum detectable glucose concentration was 100nM, which is one of the best lower
253
detection limits reported [14-19]. A large cathodic current is observed in presence of glucose
254
and the catalytic current decreases with decrease in concentration of glucose. Glucose reacts
255
with GOx in presence of oxygen to produce gluconic acid and H2O2. The H2O2 formed during
256
the reaction acted as a substrate for the second enzyme, HRP and we measured HRP
257
catalyzed reduction current of H2O2. A large cathodic electroreduction current of H2O2 is
258
observed in presence of glucose and the catalytic current decreases with decrease in
259
concentration of glucose. The heme which has come out or exposed from the HRP is reduced
260
to Fe2+ state in the applied potential region. This acts as a catalyst for the electroreduction of
261
H2O2 generated by glucose and converted to Fe3+.
262 263
A schematic representation of the enzymatic reaction mechanism of Au/MPA/CBUAuNP/GOx-HRP electrode is given in figure 3B.
264 265
Fig.3. (A) Cyclic voltammogram of Au/MPA/CBU-AuNP/GOx-HRP in the absence of glucose
266
and in the presence of different concentrations of glucose. In figure, (a) corresponds to the
267
cyclic voltammograms in the absence of glucose, (b-f) correspond to the cyclic
268
voltammograms of the electrode in100nM, 1µM, 10µM, 100µM and 1mM of glucose
269
respectively and (B) Schematic representation of mechanism of sensing.
270 271
Gold nanoparticles present on the electrode surface acted as tiny conducting centers and
272
mediated an efficient electron transfer between the enzyme and electrode surface possibly
11
273
through tunneling [26], enhancing the sensitivity of the sensor. In addition, the CBU-AuNP
274
acted as a favorable biocompatible environment for the active immobilization of enzymes
275
owing to their hydrogen bonding capability. Here the immobilizations of enzymes were
276
achieved through the hydrogen bonding between enzymes and CBU, which helps to retain the
277
activity of enzymes after immobilization.
278
For a comparison, a monoenzyme sensor surface, Au/MPA/CBU-AuNP/GOx was
279
fabricated and used for the detection of glucose. CV of Au/MPA/CBU-AuNP/GOx sensor is
280
given in figure 4.
281 282
Fig.4.
Cyclic
voltammogram
283
concentrations of glucose
of
Au/MPA/CBU-AuNP/GOx
electrode
in
different
284 285
It has been observed that the ldl is limited with 1µM glucose concentration. The
286
detection mechanism is different in this case. The voltammetric current response is found to
287
increase with increase in concentration of glucose. This change in current at higher potential
288
is owing to the oxidation of hydrogen peroxide generated during the glucose oxidase induced
289
oxidation of glucose. Since the oxidation requires large over potential (above 1V) the
290
detection of glucose using the immobilization of GOX alone would not be a promising
291
method. Here comes the advantage of the bienzyme electrode system which requires a
292
smaller reduction potential (less than -0.4 V) for the detection of glucose. Electrochemical
293
behavior of Au/MPA/CBU-AuNP/HRP electrode in glucose was also studied. No detectable
294
change in current was noted which supported the above said mechanism. The result is given
295
as figure S4 in supplementary information.
296 297
12
298
3.3. Chronoamperometry
299 300
Chronoamperometric measurements were carried out to evaluate the sensitivity of the
301
sensor surface. The experiment was performed with varying concentrations of glucose in
302
0.1M PBS at an applied potential of -0.3V. Sensor surface achieved 95% of the steady state
303
current within 5ms (Figure 5A). This fast response is due to the rapid electron transfer
304
between the exposed heme and the gold electrode, due to the favorable orientation of enzyme
305
on the conducting gold nanoparticles on the surface. The current decreases steeply with time
306
and an increase in current was observed with increase in concentration of glucose. Current
307
response of the electrode surface with concentration of glucose is shown in figure 5B which
308
is in consistent with the characteristics of Michaelis-Menten kinetics. The apparent
309
Michaelis-Menten constant (
310
electrochemical version of Lineweaver-Burk plot (figure 5C) according to Lineweaver-Burk
311
equation [10];
) describing the enzymatic affinity is calculated from the
312 313
(2)
314 315
where,
is the steady state current after the addition of the substrate, C is the bulk
316
concentration of the substrate and
317
substrate condition.
318
formation. Thus
319
value depends on the relative values of
320
is the maximum current measured under saturated is the ratio of the rates of breakdown of ES to its rate of
becomes a measure of the affinity of an enzyme for its substrate, since its and
for ES formation and dissociation. A low
value indicates strong substrate binding. The
value of Au/MPA/CBU-AuNP/GOx-
321
HRP electrode is 7.3µM. The low value of
indicates the high affinity of enzyme to the
322
substrate demonstrating the immobilization of enzyme in the active form. The proposed 13
323
sensor exhibited a linear range from 100nM to 1mM (figure 5C). For better clarity, the figure
324
is plotted from 1nM to 1M (this range is used in the real blood analysis) and entire range is
325
given in the supplementary information (figure S5). The linear range reported here is in the
326
similar range as reported in literature and is found to be as well within the useful range for
327
practical clinical applications (see below for the real sample analysis). The sensitivity of the
328
electrode was calculated as 217.5µAmM-1cm-2.
329 330
Fig.5. (A) Plot of catalytic current vs. Concentration of glucose. Inset shows the
331
chronoamperometric response of Au/MPA/CBU-AuNP/GOx-HRP electrode in different
332
concentrations of glucose, (B) Lineweaver- Burk plot of 1/Iss vs. 1/Concentration of glucose
333
and (C) Linear calibration curve of current vs. Concentration of glucose
334 335
3.4. Determination of kinetic parameters
336 337
The apparent rate of catalysis or the turnover rate constant, k'cat which represents the
338
maximum number of substrate molecules that can be converted into products per catalytic
339
site, is calculated from the Lineweaver-Burk plot according to the equation [29],
340 341
(3)
342 343
where,
is the maximum current, i.e., when the enzyme activity is saturated by
344
substrate, n is the number of electrons transferred, F is the Faraday constant and A is the area
345
of the electrode. The calculated k'cat for the sensor surfaces with various ratios of the enzymes
346
are presented in table 2. It is clear that the highest k'cat is observed for the 1:1 ratio of the
347
enzymes on the surface and hence is the best performing surface.
14
348
Biosensor performance is usually measured in terms of its catalytic efficiency,
349
where k'cat is the apparent rate of catalysis and
350
the value better is the sensor performance.
,
is the Michaelis-Menten constant. Higher
351
We compared the apparent rate of catalysis and efficiency of bienzyme glucose sensor in
352
different concentration ratios of GOx and HRP. The sensor using a 1:1 molar ratio of
353
GOx/HRP showed highest rate of catalysis and efficiency. The experimental observation is
354
further validated from the values of Gibbs activation energy. For an enzyme catalyzed
355
reaction, the Gibbs activation energy
is [29],
356 357
(4)
358 359
Where, R, T, h and
are the universal gas constant, temperature, Planck’s constant
360
and Boltzmann’s constant respectively. The Gibbs activation energy - the energy between the
361
transition state and ground state of the enzymatic reaction- was calculated for the sensor
362
surfaces with different molar ratios of enzymes viz. 1:1, 1:2, 1:4 and 1:8 of GOx and HRP.
363
The sensor surface modified with an enzyme concentration ratio of 1:1 showed lowest
364
activation energy of 76.09kJ/mol. The kinetic constants of different systems have been
365
tabulated in table 2.
366 367
Table. 2. Kinetic constants and catalytic efficiencies of the sensor system at various molar
368
ratios of GOx and HRP
369 370 371 372 15
373
3.5. Selectivity, reproducibility and stability
374 375
The selectivity of the biosensor was evaluated by studying the interference of
376
electroactive compounds such as bovine serum albumin (BSA), ascorbic acid (AA) and urea
377
on sensor performance. Figures 6 show the high selectivity of electrode surface to glucose. It
378
was seen that the sensor surface gives unaffected analytical responses in presence of the
379
interfering compounds.
380 381
Fig.6. Cyclic voltammetric response of the Au/MPA/CBU-AuNP/GOx-HRP sensor surface in
382
buffer, in 1mM glucose and in a mixture of 1mM AA, 1mM BSA, 1mM urea and 1mM glucose
383 384
Fabrication reproducibility of the sensor surface was assessed by preparing four
385
electrodes independently and measuring the response for 1mM glucose. All the electrodes
386
showed almost the same reduction current value. The fabrication reproducibility of four
387
sensor electrodes gave a relative standard deviation (RSD) of 4.5% for the determination of
388
1mM glucose.
389
The stability of the sensor surface was studied with its response to 1mM glucose over a
390
period of one month. The electrode, when not in use was stored at 4°C. The results indicated
391
that 85% of the response was retained after storage of one month.
392
In order to appraise the practical usage of Au/MPA/CBU-AuNP/GOx-HRP sensor
393
surface, the sensor surface was used for the determination of glucose in human blood serum,
394
whole blood and blood plasma samples. For determination of glucose in blood serum
395
samples, a 10μL of blood serum was added to 10mL of 0.1M PBS of pH 7.4 and the
396
chronoamperometric response was recorded at -0.3V. The average glucose concentration
397
determined was 79.8 mg/dl with a relative standard deviation of 1.23% which is in well
16
398
agreement with the value measured from the local health clinic (80 mg/dl) demonstrating that
399
the sensor surface has great practical usage for clinical analysis. The sensor was also used for
400
the detection of glucose in whole blood and blood plasma samples. For this, 10L of whole
401
blood or blood plasma sample was added to 10mL of 0.1M PBS of pH 7.4 and the
402
chronoamperometric response was recorded as in the previous experiment. The average
403
glucose concentration obtained is 80.6 mg/dl in the whole blood (with RSD 1.1%) and 79.8
404
mg/dl in the blood plasma sample (RSD 1.23 %) which is in close agreement with the value
405
obtained from the local health clinic (80mg/dl). All the experiments suggest the high
406
sensitivity of the sensor and its practical application in clinical analysis. A comparison of this
407
sensor performance with other bienzyme sensors reported are presented in table S2 of
408
supplementary information.
409 410
4. Conclusions
411 412
We have developed a sensitive bi-enzyme glucose sensor by the co-immobilization of
413
GOx and HRP on a CBU-AuNP modified gold electrode by exploiting the electrocatalytic
414
activity and efficient electron tunneling property of gold nanoparticles along with the
415
biocompatible environment provided by CBU. An improved detection limit of 100nM has
416
been obtained with a high sensitivity of 217.5µAmM-1cm-2 and a low
417
detailed investigation on the concentration ratios of GOx and HRP in the sensor performance
418
has been performed. The experimentally observed result was verified by calculating the
419
kinetic parameters such as
420
compared to nonenzymatic sensors, the sensitivity, selectivity and fabrication reproducibility
421
for these enzymatic sensors are higher. Moreover, the electrode surface can be reused if
, k'cat and
value of 7.3µM. A
. Even though the cost of the sensor is more
17
422
stored under the right conditions. We are envisaging the utilization of proposed strategy for
423
glucose detection in clinical laboratories.
424 425
Acknowledgements
426 427
Authors are grateful for the financial support from Department of Science and Technology,
428
India. We thank Mr. Shanmugharaja for assistance with the Contact angle measurements.
429
M.M acknowledges CSIR, India for research fellowship.
430 431
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comparison of mono-enzyme GOD and bienzyme GOD–HRP systems, Sens.
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glucose oxidase and biosensing for glucose based on boron-doped carbon
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nanotubes modified electrode, Biosens. Bioelectron. 23 (2008) 1272–1277.
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phthalocyanine–glucose oxidase biocomposite, Biosens. Bioelectron. 20 (2005) 1388-
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1396.
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[23] S. Zhang, N. Wang, Y. Niu, C. Sun, Immobilization of glucose oxidase on gold
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nanoparticles modified Au electrode for the construction of biosensor, Sens. Actuators,
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B, 109 (2005) 367-374.
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[24] D. Ranganathan, C. Lakshmi, L.K. Isabella, Hydrogen-bonded self-assembled peptide
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[25] S. Bharathi, M. Nogami, S. Ikeda, Novel electrochemical interfaces with a tunable
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510
horseradish peroxidase at carbon nanotube powder microelectrode, Sens. Actuators, B.
511
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514
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515
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516
and experimental approach, Sens. Actuators, B. 122 (2007) 395–402.
517 518
[29] E.I. Iwuoha, M.R. Smyth, M.E.G. Lyons, Solvent effects on the reactivities of an amperometric glucose sensor, J. Electroanal. Chem. 390 (1995) 35-45.
519
22
Figure Captions Figure 1. Schematic representation of Au/MPA/CBU-AuNP/GOx-HRP sensor surface on each stages of fabrication (A)MPA monolayer on Au (B)Assembled CBU-AuNP on MPA monolayer (C)Immobilised enzymes on the CBU-AuNP surface. Figure 2. Cyclic voltammograms of
bare Au,
Au/MPA, Au/MPA/CBU-AuNP and
Au/MPA/CBU-AuNP/GOx-HRP electrodes in 0.1M KCl solution containing 5mM K3[Fe(CN)6] Figure 3. (A) Cyclic voltammogram of Au/MPA/CBU-AuNP/GOx-HRP in the absence of glucose and in the presence of different concentrations of glucose. In figure, (a) corresponds to the cyclic voltammograms in the absence of glucose, (b-f) correspond to the cyclic voltammograms of the electrode in100nM, 1µM, 10µM, 100µM and 1mM of glucose respectively and (B) Schematic representation of mechanism of sensing. Figure 4. Cyclic voltammograms of Au/MPA/CBU-AuNP/GOx electrode in different concentrations of glucose Figure 5. (A) Plot of catalytic current vs. Concentration of glucose. Inset shows the chronoamperometric response of Au/MPA/CBU-AuNP/GOx-HRP electrode in different concentrations of glucose, (B) Lineweaver- Burk plot of 1/Iss vs. 1/Concentration of glucose and (C) Linear calibration curve of current vs. concentration of glucose Figure 6. Cyclic voltammetric response of the Au/MPA/CBU-AuNP/GOx-HRP sensor surface in buffer, in 1mM glucose and in a mixture of 1mM AA, 1mM BSA, 1mM urea and 1mM glucose
Figures Figure1.
Figure 2.
Figure 3.
Figure 4.
Figure 5.
Figure 6.
List of Tables
Table.1. Quartz crystal frequency change and the mass adsorbed on the crystal surface after the immobiliztion of MPA, CBU-AuNP, GOx-HRP
Monolayer
Resonance Frequency
Frequency shift (Hz)
Mass adsorbed (ng)
(Hz) Au
7979281
-
-
Au/MPA
7979256
25
34.25
Au/MPA/CBU-AuNP
7979111
145
210.55
Au/MPA/CBU-
7979022
89
123.71
AuNP/GOx-HRP
Table. 2. Kinetic constants and catalytic efficiencies of the sensor system at various molar ratios of GOx and HRP
Enzyme
k'cat
′
# (kJmol-1)
-2 -1
concentration ratios (pmolcm s )
(µM)
(cms-1)
[GOx]:[HRP] 1:1
0.219
7.3
0.291
76.09
1:2
0.093
8.91
0.105
78.59
1:4
0.036
15.68
0.018
82.89
1:8
0.023
20.16
0.011
84.07
S0003-2697(14)00198-5 http://dx.doi.org/10.1016/j.ab.2014.05.003 YABIO 11736
To appear in:
Analytical Biochemistry
Received Date: Revised Date: Accepted Date:
10 February 2014 29 April 2014 6 May 2014
Please cite this article as: M. Mathew, N. Sandhyarani, Detection of glucose using immobilised bi-enzyme on cyclic bisureas-gold nanoparticle conjugate, Analytical Biochemistry (2014), doi: http://dx.doi.org/10.1016/j.ab. 2014.05.003
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1
Detection of glucose using immobilised bi-enzyme on cyclic
2
bisureas-gold nanoparticle conjugate
3 4 5 6 7 8
Manjusha Mathew, N. Sandhyarani* Nanoscience Research Laboratory, School of Nano Science and Technology, National Institute of Technology Calicut, Calicut, Kerala, India *Author to whom correspondence to be addressed. e-mail: [email protected] Ph: 91-495 2286537, Fax: 91-495 2287250
9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24
1
25
Abstract
26 27
A highly sensitive electrochemical glucose sensor has been developed by the co-
28
immobilization of glucose oxidase (GOx) and horseradish peroxidase (HRP) onto a gold
29
electrode modified with biocompatible cyclic bisureas-gold nanoparticle conjugate (CBU-
30
AuNP). A self-assembled monolayer of mercapto propionic acid (MPA) and CBU-AuNP was
31
formed on the gold electrode through a layer by layer assembly. This modified electrode was
32
used for immobilisation of the enzymes GOx and HRP. Both the HRP and GOx retained their
33
catalytic activity for an extended time indicated by the low value of Michaelis-Menten
34
constant. Analytical performance of the sensor was examined in terms of sensitivity,
35
selectivity, reproducibility, lower detection limit and stability. The developed sensor surface
36
exhibited a limit of detection of 100nM with a linear range of 100nM to 1mM. A high
37
sensitivity of 217.5µAmM-1cm-2 at a low potential of -0.3V was obtained in this sensor
38
design. Various kinetic parameters were calculated. The sensor was examined for its practical
39
clinical application by estimating glucose in human blood sample.
40 41 42
Keywords- glucose sensor; enzyme; glucose oxidase; horseradish peroxidase; cyclic
43
voltammetry
44 45 46 47 48 49
2
50
1. Introduction
51 52
Accurate determination and regular monitoring of blood glucose concentration is very
53
important in the diagnosis and treatment of diabetes mellitus. In addition, sensitive and
54
accurate detection of glucose has extensive applications in food industry and biotechnology
55
[1,2]. Specific recognition of glucose by glucose oxidase and reactions of glucose by glucose
56
oxidase are widely explored in commonly used electrochemical glucose sensors [3-5]. Since
57
the flavine adenine dinucleotide (FAD) redox centre of GOx is deeply buried inside a
58
protective protein shell, it is difficult to establish a direct electron communication between
59
the enzyme and electrode. In order to improve the sensitivity, mediated electron transfer
60
using ferrocene derivatives [6,7] and electroactive polymers [8] were reported. Application of
61
mediators provides higher sensitivity, however their application is limited as they
62
contaminate the sample solution. Consequently, modification of electrode surface to allow
63
efficient electron transfer between the active site of redox protein and electrode become
64
significant in the development of biosensors. Formation of well-ordered, homogeneous and
65
molecular monolayers on the surface of electrode by self assembly [9] can facilitate electron
66
transfer between the analyte and the electrode, offering good reproducibility with a fast
67
response [10]. In addition, integration of nanostructured materials with enzymatic biosensor
68
is widely used to improve the electron transfer between the GOx and electrode [11-13].
69
Enzymes are highly sensitive and may undergo denaturation with variation in conditions like
70
temperature, pH, humidity etc. Hence, enzyme attachment to the solid substrate and its
71
compatibility with the metal electrode is crucial in the realization of a sensitive biosensor.
72
Consequently, a biocompatible environment is needed to sustain the enzyme activity for an
73
extended time.
3
74
Co-immobilization of HRP with a H2O2 producing oxidase enzyme to fabricate a bi-
75
enzyme biosensor is an effective way to improve the glucose sensor performance [14-19]. In
76
this system, the H2O2 produced as a result of the enzyme induced oxidation of glucose is
77
subsequently reduced by HRP which permits the detection of glucose at a low potential. Thus
78
the bi-enzyme system switches the detection principle from an electrochemical oxidation
79
reaction to reduction reaction.
80
We have reported the efficient use of CBU-AuNP for the development of a hydrogen
81
peroxide sensor. Owing to the biocompatible environment of CBU and electron transfer
82
capability of gold nanoparticles we achieved one of the lowest reported limits of detection for
83
hydrogen peroxide [20]. The biocompatible environment and efficient electron transfer
84
ability provided by the CBU-AuNP have also been employed in the development of a glucose
85
sensor. On using the immobilized GOx on the surface we could detect the glucose only up to
86
1 M which is in the similar range as that of most of the GOx based glucose sensors [21-23].
87
To improve the detection limit of glucose sensor we have used a bi-enzyme strategy and
88
achieved a nanomolar level detection limit for glucose.The work reported in this manuscript
89
demonstrates the development of a highly sensitive glucose sensor using the immobilization
90
of GOx and HRP- on a gold electrode modified with CBU capped gold nanoparticles. H2O2
91
produced as a result of glucose oxidation by GOx served as a substrate for the HRP.
92
Optimized conditions of sensor performance were studied. Selectivity, reproducibility and
93
stability of the sensor surface were also evaluated and the sensor was used for the
94
measurement of the glucose level in blood.
95
2. Experimental Methods
96
2.1. Materials
97 98
GOx from Aspergillus niger, Peroxidase from horseradish type V1-A, hydrogen tetrachloroaurate trihydrate, 1,6-diisocyanato hexane and L-cystine
4
dimethyl ester
99
dihydrochloride were
obtained from Sigma-Aldrich. Monosodium glutamate, dextrose
100
anhydrous, Na2CO3, Na2HPO4 and KH2PO4 were purchased from Merck international. The
101
supporting electrolyte 0.1M Phosphate Buffered Saline buffer was prepared by mixing the
102
stock solutions of 0.1M Na2HPO4 and KH2PO4 and the pH was adjusted using NaOH or HCl.
103
All solutions were prepared with ultrapure water obtained from an ultrafiltration system
104
(18M cm. Milli-Q, Millipore system).
105 106
2.2. Synthesis
107 108
A standard procedure was adopted for the synthesis of CBU [24]. The hydrogen
109
bonding possibility of CBU has been employed for the immobilization of enzymes. To
110
achieve efficient electron transport between the gold electrode and analyte the CBU was
111
conjugated to gold nanoparticles. CBU-AuNP was prepared as described elsewhere [20].
112 113
2.3. Preparation of electrodes
114 115
Gold electrode of geometric area 0.28cm2 was mechanically polished with 0.3μm -
116
Al2O3 slurry and washed with distilled water. Then the electrode was sonicated in a freshly
117
prepared Piranha solution for 10min. It was then washed with deionised water, sonicated and
118
rinsed with ethanol and dried. A monolayer of MPA was prepared by immersing the cleaned
119
gold electrode in 1mM MPA in ethanol for 12h at room temperature. The electrode was taken
120
out and thoroughly rinsed with ethanol to remove the physically adsorbed thiol and dried.
121
Then the electrode was placed into CBU-AuNP composite in methanol for 4h. Finally
122
modification by the enzymes to the electrode surface was attained by incubating the electrode
5
123
in a mixture of 0.1mM GOx and 0.1mM HRP in PBS for 12h. After thoroughly rinsing the
124
sensor with deionised water, it was kept at 4°C when not in use.
125 126
2.4. Characterizations
127 128
Cyclic voltammetry (CV), chronoamperometry and quartz crystal microbalance -
129
QCM (Resonance frequency of the crystal - 7.9MHz) measurements were performed using
130
CH 400A Electrochemical Quartz Crystal Microbalance (CH Instruments, Austin, Texas). A
131
three electrode system was used with the modified gold electrode as working electrode, a
132
platinum wire as auxiliary electrode and saturated calomel as the reference electrode against
133
which all potentials were measured. The measurements have been carried out at room
134
temperature with a scan rate of 100mV/s. QCM experiments were carried out on a gold
135
polished quartz crystal electrode as a working electrode, calomel electrode as reference
136
electrode and a platinum wire as counter electrode. Atomic force microscopy (AFM)
137
experiments were performed using Park XE-100 atomic force microscope. Typical AFM
138
images were acquired in non contact mode with silicon nitride cantilever. Energy dispersive
139
X-ray spectroscopy (EDS) measurements were performed using
140
pressure field emission scanning electron microscope at an accelerating voltage 15kV.
141
Contact angle measurements were performed with a stereo zoom microscope. A micropipette
142
was used to dispense 10μL of millipore water on electrode surface and placed under the
143
objective. All measurements were carried out at room temperature.
144 145 146 147
6
Hitachi SU6600 variable
148
3.
Results and discussion
149 150
A schematic representation of the sensor surface is shown in figure 1. Initially the
151
MPA monolayers were formed on the cleaned gold electrode which was then used for the
152
assembly of CBU-AuNP. On this layer, GOx and HRP have been immobilised by taking
153
advantage of the large hydrogen bonding possibility of CBU.
154 155
Fig.1. Schematic representation of Au/MPA/CBU-AuNP/GOx-HRP sensor surface on each
156
stages of fabrication (A) MPA monolayer on Au (B)Assembled CBU-AuNP on MPA
157
monolayer (C)Immobilised enzymes on the CBU-AuNP surface.
158 159
3.1. Characterization of electrode surface
160 161
3.1.1. Cyclic voltammetry
162 163
Formation of each layer on the electrode surface was monitored using the cyclic
164
voltammogram of ferro/ferricyanide redox couple. The experiment was conducted in the
165
electrolyte containing 5mM K3[Fe(CN)6] and 0.1M KCl at a scan rate of 100mVs-1 for four
166
different types of electrodes viz. bare Au, Au/MPA, Au/MPA/CBU-AuNP and
167
Au/MPA/CBU-AuNP/GOx-HRP. Figure 2 compares the voltammetric responses at each
168
stages of the fabrication process.
169 170
Fig.2. Cyclic voltammograms of
bare Au,
171
Au/MPA/CBU-AuNP/GOx-HRP electrodes in 0.1M KCl solution containing 5mM
172
K3[Fe(CN)6].
7
Au/MPA, Au/MPA/CBU-AuNP and
173 174
A well defined redox peak of the Fe2+/Fe3+ was observed on bare gold electrode.
175
Modification of the electrode surface with MPA causes a decrease in the anodic and cathodic
176
current as it forms an insulating layer on the electrode, which acted as a blockade to the
177
electron transfer between the analyte and electrode surface. A remarkable increase of current
178
is obtained after CBU-AuNP assembly due to the electron transfer through the gold
179
nanoparticles. In addition to the electron transfer property, the high surface area of gold
180
nanoparticles and its enhanced catalytic activity [25,26] also plays an important role in
181
enhancing the current. Immobilization of enzymes on the electrode decreases the redox
182
current values due to the insulating nature of the enzymes.
183 184
3.1.2. Electrochemical quartz crystal microbalance
185 186
The surface modification of electrode by various immobilization steps are also confirmed
187
by monitoring the mass change in electrochemical quartz crystal microbalance (EQCM). The
188
change in resonance frequency of the crystal due to the increase in mass as a result of
189
adsorption of different materials on the electrode surface was recorded. The relationship
190
between the changes in mass per unit area (Δm) and frequency (Δf) are given by Sauerbrey
191
equation [27];
192 193
(1) where,
,
,
,
,
and
represent the frequency change, resonance frequency of
194
the fundamental mode(Hz), mass change(ng), area of the crystal (0.205cm2), shear
195
modulus(g/cms2) and density(g/cm3) of the crystal respectively.
196
EQCM results indicate that formation of different layers resulted in a decrease in
197
frequency and an increase in mass on the surface. The immobilization of MPA induces a shift 8
198
in the frequency of 25Hz, which corresponds to the adsorbed mass of 34.75 ng. The surface
199
concentration of MPA is calculated as 9.6×1014 molecules/cm2. Immobilization of CBU-
200
AuNP leads to a deposition of 201.55ng on the quartz crystal. Deposition of enzymes resulted
201
in a frequency change of 89Hz corresponding to a deposition of 123.71ng of enzymes.
202
(Molecular weight of GOx is 160kD and HRP is 44kD). The surface concentration of enzyme
203
is obtained as 1.7×1012 molecules/cm2 assuming equal probability of immobilization of both
204
the enzymes leading to equal concentration of enzymes on the surface. Table 1 summarizes
205
the quartz crystal frequency change and the mass adsorbed on the crystal surface after each
206
steps of immobilization.
207 208
Table.1. Quartz crystal frequency change and the mass adsorbed on the crystal surface after
209
the immobiliztion of MPA, CBU-AuNP, GOx-HRP
210 211
EDS is an effective tool to probe the change in weight percentage of elements on each
212
stages of immobilization. Table S1 of supplementary information summarizes the weight
213
percentage obtained on successive immobilization of different layers on the electrode surface.
214
Increases in weight percentages of carbon, sulphur and nitrogen confirm the successful
215
immobilization of each layer on the electrode surface.
216
Contact angle of the sensor surface on each stages of immobilization was also
217
measured. Microscopic images of a water droplet placed on each layer of Au/MPA/CBU-
218
AuNP/GOx-HRP electrode are given in figure S1 of supplementary information. A bare gold
219
electrode showed a contact angle of 72.8°. A decrease in contact angle to 62° is observed on
220
immobilization of MPA indicating an increased wetting due to the terminating carboxyl
221
group. Presence of biocompatible gold nanoparticles further increased the hydrophilicity of
222
the surface leading to a decrease in the contact angle value to 56°. Immobilization of enzymes
9
223
on the electrode leads to an additional reduction in the contact angle to 43.6°. The increase in
224
hydrophilic nature confirms the presence of enzymes on the electrode surface. The increased
225
hydrophilicity of gold nanoparticles helps to retain the activity of enzyme on the electrode
226
surface. Atomic force microscopic images also were taken to confirm the immobilization of
227
each step which is presented in supplementary information (figure S2).
228 229
3.2. Electrocatalytic oxidation of glucose at the bi-enzyme electrode
230 231
Optimum response of the electrode surface under different concentrations of the enzymes
232
were investigated and found that the molar ratio used for immobilization was critical in
233
determining the experimentally observed lower detection limit (ldl) [28]. Optimum
234
concentrations of [GOx]/[HRP] were determined by measuring the sensor performance by
235
varying the molar ratios of enzymes used for immobilisation. The current response and lower
236
detection limit in each case were determined. Three different molar ratios of GOx and HRP
237
viz. 1:1, 2:1 and 1:2 were selected for the preliminary experiments and the observed ldl was
238
of 100nM in the case of 1:1 and 1:2 and 1μM in the case of 2:1 ratio of GOx and HRP. Figure
239
S3A shows a comparison of the limit of detection of the three sensor surfaces. Since the 1:1
240
and 1:2 ratios of GOx and HRP exhibited the same ldl, we varied the concentration of HRP
241
from 1 to 8 times keeping the GOx concentration constant and determined the ldl. The lowest
242
detection limit was achieved with molar ratios of 1:1 and 1:2 of GOx and HRP respectively.
243
The result is given in the supplementary information as figure S3B. The current response
244
obtained for the oxidation of 1mM glucose with different molar ratios of [GOx]/[HRP] were
245
monitored by the change in current of H2O2 reduction. Current values of 90µA, 50µA, 49µA
246
and 35µA were obtained for the GOx:HRP concentration ratios of 1:1, 1:2, 1:4 and 1:8
247
respectively. Since the electrode prepared with GOx and HRP at a molar ratio of 1:1 yielded
10
248
maximum current response for 1mM glucose, we selected 1:1 molar ratio of GOx and HRP as
249
the optimum concentration and the response of the resulted electrode was studied in detail.
250
Figure 3A shows the response of the sensor at various concentrations of glucose. An
251
increase in cathodic current was observed with increase in concentration of glucose.
252
Minimum detectable glucose concentration was 100nM, which is one of the best lower
253
detection limits reported [14-19]. A large cathodic current is observed in presence of glucose
254
and the catalytic current decreases with decrease in concentration of glucose. Glucose reacts
255
with GOx in presence of oxygen to produce gluconic acid and H2O2. The H2O2 formed during
256
the reaction acted as a substrate for the second enzyme, HRP and we measured HRP
257
catalyzed reduction current of H2O2. A large cathodic electroreduction current of H2O2 is
258
observed in presence of glucose and the catalytic current decreases with decrease in
259
concentration of glucose. The heme which has come out or exposed from the HRP is reduced
260
to Fe2+ state in the applied potential region. This acts as a catalyst for the electroreduction of
261
H2O2 generated by glucose and converted to Fe3+.
262 263
A schematic representation of the enzymatic reaction mechanism of Au/MPA/CBUAuNP/GOx-HRP electrode is given in figure 3B.
264 265
Fig.3. (A) Cyclic voltammogram of Au/MPA/CBU-AuNP/GOx-HRP in the absence of glucose
266
and in the presence of different concentrations of glucose. In figure, (a) corresponds to the
267
cyclic voltammograms in the absence of glucose, (b-f) correspond to the cyclic
268
voltammograms of the electrode in100nM, 1µM, 10µM, 100µM and 1mM of glucose
269
respectively and (B) Schematic representation of mechanism of sensing.
270 271
Gold nanoparticles present on the electrode surface acted as tiny conducting centers and
272
mediated an efficient electron transfer between the enzyme and electrode surface possibly
11
273
through tunneling [26], enhancing the sensitivity of the sensor. In addition, the CBU-AuNP
274
acted as a favorable biocompatible environment for the active immobilization of enzymes
275
owing to their hydrogen bonding capability. Here the immobilizations of enzymes were
276
achieved through the hydrogen bonding between enzymes and CBU, which helps to retain the
277
activity of enzymes after immobilization.
278
For a comparison, a monoenzyme sensor surface, Au/MPA/CBU-AuNP/GOx was
279
fabricated and used for the detection of glucose. CV of Au/MPA/CBU-AuNP/GOx sensor is
280
given in figure 4.
281 282
Fig.4.
Cyclic
voltammogram
283
concentrations of glucose
of
Au/MPA/CBU-AuNP/GOx
electrode
in
different
284 285
It has been observed that the ldl is limited with 1µM glucose concentration. The
286
detection mechanism is different in this case. The voltammetric current response is found to
287
increase with increase in concentration of glucose. This change in current at higher potential
288
is owing to the oxidation of hydrogen peroxide generated during the glucose oxidase induced
289
oxidation of glucose. Since the oxidation requires large over potential (above 1V) the
290
detection of glucose using the immobilization of GOX alone would not be a promising
291
method. Here comes the advantage of the bienzyme electrode system which requires a
292
smaller reduction potential (less than -0.4 V) for the detection of glucose. Electrochemical
293
behavior of Au/MPA/CBU-AuNP/HRP electrode in glucose was also studied. No detectable
294
change in current was noted which supported the above said mechanism. The result is given
295
as figure S4 in supplementary information.
296 297
12
298
3.3. Chronoamperometry
299 300
Chronoamperometric measurements were carried out to evaluate the sensitivity of the
301
sensor surface. The experiment was performed with varying concentrations of glucose in
302
0.1M PBS at an applied potential of -0.3V. Sensor surface achieved 95% of the steady state
303
current within 5ms (Figure 5A). This fast response is due to the rapid electron transfer
304
between the exposed heme and the gold electrode, due to the favorable orientation of enzyme
305
on the conducting gold nanoparticles on the surface. The current decreases steeply with time
306
and an increase in current was observed with increase in concentration of glucose. Current
307
response of the electrode surface with concentration of glucose is shown in figure 5B which
308
is in consistent with the characteristics of Michaelis-Menten kinetics. The apparent
309
Michaelis-Menten constant (
310
electrochemical version of Lineweaver-Burk plot (figure 5C) according to Lineweaver-Burk
311
equation [10];
) describing the enzymatic affinity is calculated from the
312 313
(2)
314 315
where,
is the steady state current after the addition of the substrate, C is the bulk
316
concentration of the substrate and
317
substrate condition.
318
formation. Thus
319
value depends on the relative values of
320
is the maximum current measured under saturated is the ratio of the rates of breakdown of ES to its rate of
becomes a measure of the affinity of an enzyme for its substrate, since its and
for ES formation and dissociation. A low
value indicates strong substrate binding. The
value of Au/MPA/CBU-AuNP/GOx-
321
HRP electrode is 7.3µM. The low value of
indicates the high affinity of enzyme to the
322
substrate demonstrating the immobilization of enzyme in the active form. The proposed 13
323
sensor exhibited a linear range from 100nM to 1mM (figure 5C). For better clarity, the figure
324
is plotted from 1nM to 1M (this range is used in the real blood analysis) and entire range is
325
given in the supplementary information (figure S5). The linear range reported here is in the
326
similar range as reported in literature and is found to be as well within the useful range for
327
practical clinical applications (see below for the real sample analysis). The sensitivity of the
328
electrode was calculated as 217.5µAmM-1cm-2.
329 330
Fig.5. (A) Plot of catalytic current vs. Concentration of glucose. Inset shows the
331
chronoamperometric response of Au/MPA/CBU-AuNP/GOx-HRP electrode in different
332
concentrations of glucose, (B) Lineweaver- Burk plot of 1/Iss vs. 1/Concentration of glucose
333
and (C) Linear calibration curve of current vs. Concentration of glucose
334 335
3.4. Determination of kinetic parameters
336 337
The apparent rate of catalysis or the turnover rate constant, k'cat which represents the
338
maximum number of substrate molecules that can be converted into products per catalytic
339
site, is calculated from the Lineweaver-Burk plot according to the equation [29],
340 341
(3)
342 343
where,
is the maximum current, i.e., when the enzyme activity is saturated by
344
substrate, n is the number of electrons transferred, F is the Faraday constant and A is the area
345
of the electrode. The calculated k'cat for the sensor surfaces with various ratios of the enzymes
346
are presented in table 2. It is clear that the highest k'cat is observed for the 1:1 ratio of the
347
enzymes on the surface and hence is the best performing surface.
14
348
Biosensor performance is usually measured in terms of its catalytic efficiency,
349
where k'cat is the apparent rate of catalysis and
350
the value better is the sensor performance.
,
is the Michaelis-Menten constant. Higher
351
We compared the apparent rate of catalysis and efficiency of bienzyme glucose sensor in
352
different concentration ratios of GOx and HRP. The sensor using a 1:1 molar ratio of
353
GOx/HRP showed highest rate of catalysis and efficiency. The experimental observation is
354
further validated from the values of Gibbs activation energy. For an enzyme catalyzed
355
reaction, the Gibbs activation energy
is [29],
356 357
(4)
358 359
Where, R, T, h and
are the universal gas constant, temperature, Planck’s constant
360
and Boltzmann’s constant respectively. The Gibbs activation energy - the energy between the
361
transition state and ground state of the enzymatic reaction- was calculated for the sensor
362
surfaces with different molar ratios of enzymes viz. 1:1, 1:2, 1:4 and 1:8 of GOx and HRP.
363
The sensor surface modified with an enzyme concentration ratio of 1:1 showed lowest
364
activation energy of 76.09kJ/mol. The kinetic constants of different systems have been
365
tabulated in table 2.
366 367
Table. 2. Kinetic constants and catalytic efficiencies of the sensor system at various molar
368
ratios of GOx and HRP
369 370 371 372 15
373
3.5. Selectivity, reproducibility and stability
374 375
The selectivity of the biosensor was evaluated by studying the interference of
376
electroactive compounds such as bovine serum albumin (BSA), ascorbic acid (AA) and urea
377
on sensor performance. Figures 6 show the high selectivity of electrode surface to glucose. It
378
was seen that the sensor surface gives unaffected analytical responses in presence of the
379
interfering compounds.
380 381
Fig.6. Cyclic voltammetric response of the Au/MPA/CBU-AuNP/GOx-HRP sensor surface in
382
buffer, in 1mM glucose and in a mixture of 1mM AA, 1mM BSA, 1mM urea and 1mM glucose
383 384
Fabrication reproducibility of the sensor surface was assessed by preparing four
385
electrodes independently and measuring the response for 1mM glucose. All the electrodes
386
showed almost the same reduction current value. The fabrication reproducibility of four
387
sensor electrodes gave a relative standard deviation (RSD) of 4.5% for the determination of
388
1mM glucose.
389
The stability of the sensor surface was studied with its response to 1mM glucose over a
390
period of one month. The electrode, when not in use was stored at 4°C. The results indicated
391
that 85% of the response was retained after storage of one month.
392
In order to appraise the practical usage of Au/MPA/CBU-AuNP/GOx-HRP sensor
393
surface, the sensor surface was used for the determination of glucose in human blood serum,
394
whole blood and blood plasma samples. For determination of glucose in blood serum
395
samples, a 10μL of blood serum was added to 10mL of 0.1M PBS of pH 7.4 and the
396
chronoamperometric response was recorded at -0.3V. The average glucose concentration
397
determined was 79.8 mg/dl with a relative standard deviation of 1.23% which is in well
16
398
agreement with the value measured from the local health clinic (80 mg/dl) demonstrating that
399
the sensor surface has great practical usage for clinical analysis. The sensor was also used for
400
the detection of glucose in whole blood and blood plasma samples. For this, 10L of whole
401
blood or blood plasma sample was added to 10mL of 0.1M PBS of pH 7.4 and the
402
chronoamperometric response was recorded as in the previous experiment. The average
403
glucose concentration obtained is 80.6 mg/dl in the whole blood (with RSD 1.1%) and 79.8
404
mg/dl in the blood plasma sample (RSD 1.23 %) which is in close agreement with the value
405
obtained from the local health clinic (80mg/dl). All the experiments suggest the high
406
sensitivity of the sensor and its practical application in clinical analysis. A comparison of this
407
sensor performance with other bienzyme sensors reported are presented in table S2 of
408
supplementary information.
409 410
4. Conclusions
411 412
We have developed a sensitive bi-enzyme glucose sensor by the co-immobilization of
413
GOx and HRP on a CBU-AuNP modified gold electrode by exploiting the electrocatalytic
414
activity and efficient electron tunneling property of gold nanoparticles along with the
415
biocompatible environment provided by CBU. An improved detection limit of 100nM has
416
been obtained with a high sensitivity of 217.5µAmM-1cm-2 and a low
417
detailed investigation on the concentration ratios of GOx and HRP in the sensor performance
418
has been performed. The experimentally observed result was verified by calculating the
419
kinetic parameters such as
420
compared to nonenzymatic sensors, the sensitivity, selectivity and fabrication reproducibility
421
for these enzymatic sensors are higher. Moreover, the electrode surface can be reused if
, k'cat and
value of 7.3µM. A
. Even though the cost of the sensor is more
17
422
stored under the right conditions. We are envisaging the utilization of proposed strategy for
423
glucose detection in clinical laboratories.
424 425
Acknowledgements
426 427
Authors are grateful for the financial support from Department of Science and Technology,
428
India. We thank Mr. Shanmugharaja for assistance with the Contact angle measurements.
429
M.M acknowledges CSIR, India for research fellowship.
430 431
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22
Figure Captions Figure 1. Schematic representation of Au/MPA/CBU-AuNP/GOx-HRP sensor surface on each stages of fabrication (A)MPA monolayer on Au (B)Assembled CBU-AuNP on MPA monolayer (C)Immobilised enzymes on the CBU-AuNP surface. Figure 2. Cyclic voltammograms of
bare Au,
Au/MPA, Au/MPA/CBU-AuNP and
Au/MPA/CBU-AuNP/GOx-HRP electrodes in 0.1M KCl solution containing 5mM K3[Fe(CN)6] Figure 3. (A) Cyclic voltammogram of Au/MPA/CBU-AuNP/GOx-HRP in the absence of glucose and in the presence of different concentrations of glucose. In figure, (a) corresponds to the cyclic voltammograms in the absence of glucose, (b-f) correspond to the cyclic voltammograms of the electrode in100nM, 1µM, 10µM, 100µM and 1mM of glucose respectively and (B) Schematic representation of mechanism of sensing. Figure 4. Cyclic voltammograms of Au/MPA/CBU-AuNP/GOx electrode in different concentrations of glucose Figure 5. (A) Plot of catalytic current vs. Concentration of glucose. Inset shows the chronoamperometric response of Au/MPA/CBU-AuNP/GOx-HRP electrode in different concentrations of glucose, (B) Lineweaver- Burk plot of 1/Iss vs. 1/Concentration of glucose and (C) Linear calibration curve of current vs. concentration of glucose Figure 6. Cyclic voltammetric response of the Au/MPA/CBU-AuNP/GOx-HRP sensor surface in buffer, in 1mM glucose and in a mixture of 1mM AA, 1mM BSA, 1mM urea and 1mM glucose
Figures Figure1.
Figure 2.
Figure 3.
Figure 4.
Figure 5.
Figure 6.
List of Tables
Table.1. Quartz crystal frequency change and the mass adsorbed on the crystal surface after the immobiliztion of MPA, CBU-AuNP, GOx-HRP
Monolayer
Resonance Frequency
Frequency shift (Hz)
Mass adsorbed (ng)
(Hz) Au
7979281
-
-
Au/MPA
7979256
25
34.25
Au/MPA/CBU-AuNP
7979111
145
210.55
Au/MPA/CBU-
7979022
89
123.71
AuNP/GOx-HRP
Table. 2. Kinetic constants and catalytic efficiencies of the sensor system at various molar ratios of GOx and HRP
Enzyme
k'cat
′
# (kJmol-1)
-2 -1
concentration ratios (pmolcm s )
(µM)
(cms-1)
[GOx]:[HRP] 1:1
0.219
7.3
0.291
76.09
1:2
0.093
8.91
0.105
78.59
1:4
0.036
15.68
0.018
82.89
1:8
0.023
20.16
0.011
84.07
