Polymerization and biosensor application of water soluble peptide-SNS type monomer conjugates

Article abstract of DOI:10.1039/c7tb01674c

A simple and efficient approach for the preparation of a biosensing platform was developed based on newly designed peptide-SNS type monomer conjugates. The approach involves the electrochemical polymerization of the peptide-SNS type monomer on the electrode surface. To synthesize the peptide bearing monomers, the SNS-type monomer having a carboxylic acid functional group was anchored to the C-terminal of the peptide by solid phase peptide synthesis via coupling reagents. Utilization of peptides to increase the solubility of the monomers was first investigated in this report. The obtained monomers, soluble in water, were fully characterized by spectral analyses and utilized as matrices for biomolecule attachment. Polymerization of monomers in water has the potential to provide an alternative process for the electrochemical preparation of the polymers in aqueous media, without using any organic solvent. Under the optimized conditions, the biosensor responded to the target analyte, glucose, in a strikingly selective and sensitive manner, and showed promising feasibility for the quantitative analysis of glucose in beverages.

Full text of DOI:10.1039/c7tb01674c

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PAPER
Guoping Chen et al.
Regulating the stemness of mesenchymal stem cells by tuning
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Polymerization and biosensor application of water soluble
peptide-SNS type monomer conjugates
Received 00th January 20xx,
Accepted 00th January 20xx
a
a
a
b
a,c*
Saniye Soylemez , Tuğçe Yılmaz , Ece Buber , Yasemin A. Udum , Salih Özçubukçu and Levent
a,c,d,e*
Toppare
DOI: 10.1039/x0xx00000x
A simple and efficient approach for the preparation of biosensing platform was developed based on newly designed
peptide-SNS type monomer conjugates. The approach involves the electrochemical polymerization of the peptide-SNS
type monomer on the electrode surface. To synthesize peptide bearing monomers, the SNS-type monomer having a
carboxylic acid functional group was anchored to the C-terminal of peptide by solid phase peptide synthesis via coupling
reagents. Utilization of peptides to increase the solubility of the monomers was first investigated with this report. The
obtained monomers, soluble in water, were fully characterized by spectral analyses and utilized as matrices for
biomolecule attachment. Polymerization of monomers in water has the potential to provide an alternative process for the
electrochemical preparation of the polymers in aqueous medium, without using any organic solvent. Under the optimized
conditions, the biosensor responded to the target analyte; glucose, with a strikingly selective and sensitive manner, and
showed promising feasibility for the quantitative analysis of glucose in beverages.
www.rsc.org/
Introduction
In the past few decades increasing number of people suffered
from diabetes mellitus has made it a leading threat for human
health. The research on accurate and sensitive detection of
glucose is of great importance for early diagnosis and
detection of diabetes. Biosensors are promising devices for
glucose detection due to their high sensitivity, selectivity, short
analysis time and low cost compare to HPLC or
Peptides are an important class of molecules which cover
organic and biochemical materials and as a consequence of
this, they have a large variety of applications in biomaterials ,
7
8
9
10
medicinal chemistry , drug delivery systems , biosensors ,
etc. It is
1
organic electronics , molecular recognitions
1
12
possible to attain different structures (alpha-helices, beta
sheet etc.) and properties (hydrophilic, hydrophobic surfaces)
just by proper selection of the amino acids. They also play a
critical role in the generation of hybrid materials such as
1
spectrophotometric equipments. Due to their considerable
advantages of great electrical, electronic and optical
properties, conjugated polymers (CPs) provide a feasible
alternative in designing numerous types of biosensing
1
3
peptide-inorganic
systems
and
peptide-polymer
1
4
conjugates. Peptides with polar side chains such as aspartic
acid, glutamic acid, and arginine are mostly soluble in water.
Conjugation of the peptide having these amino acids to a
poorly soluble hydrophobic molecule would enhance its
2
,3
platforms. In addition to providing a suitable immobilization
platform for biomolecules, CPs have a tendency to enhance
2
,4-6
stability and sensitivity of the biosensors.
The requirement
1
5
for successful enzyme loading with excessive activity motivates
research into the improvement of matrices with durable and
easy bioconjugation. Functionalized CPs can facilitate covalent
immobilization which allows further modification with
materials to improve biosensor performance. Since the choice
of a suitable matrix is very important for successful detection
of analyte, adapting the electrode surface to construct such a
functional biosensor remains a challenge.
solubility in water. Conjugation of hydrophilic groups or even
polymers such as PEG to a hydrophobic monomer to increase
its water solubility is a common strategy especially in the
preparation of water soluble conjugated polymers for
1
6
biosensor applications. Same strategy was also used to
solubilize drug molecules which are very hydrophobic and
1
7
poorly soluble in water. Peptide-polymer conjugates are new
classes of materials that are applicable to wide range of areas
utilizing the features of both components to generate unique
1
8,19
properties.
biomaterials mimicking natural proteins
Moreover,
peptides
are fascinating
and they can show
2
0,21
a.
b.
c.
Department of Chemistry, Middle East Technical University, Ankara 06800, Turkey
Department of Advanced Technologies, Gazi University, Ankara 06500, Turkey
Department of Biotechnology, Middle East Technical University, Ankara 06800,
Turkey
great enhancement in the solubility of the hydrophobic
molecules in water and polar organic solvents. However,
utilization of peptides to increase the solubility of the
d.
e.
Department of Polymer Science and Technology, Middle East Technical University,
Ankara 06800, Turkey
monomers has not been investigated to the best of our
knowledge. Peptide conjugation to monomers
22,23
The Center for Solar Energy Research and Application (GUNAM), Middle East
Technical University, Ankara 06800, Turkey
can also
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Scheme 1. Schematic representation of proposed glucose biosensors.
exhibit a good platform for the immobilization of the enzymes
when the monomer is polymerized. For a biosensor design,
electrogenerated film matrices provide good stability and
2
4
improvements for biomolecule immobilization. Furthermore,
peptides on the polymer surface and enzyme molecules would
show versatile interactions since they are both made of amino
acids. This phenomenon may allow producing high
performance biosensors and long term storage in immobilized
surfaces.
Taking into consideration of the above-mentioned views,
herein we report the synthesis, characterization, and
biosensing application of water soluble SNS-type monomers
bearing peptide side chains. Although there are several
methods to conjugate peptides to polymers, covalent
attachment is one of the most common ways to anchor
peptides to polymers. However, it is also possible to conjugate
Experimental
2
5
peptides to monomers during solid phase peptide synthesis.
Materials
Hence in this work, firstly a SNS-type carboxylic group
containing monomer was synthesized and then, two different
peptides (sequences of glutamic acid-arginine-arginine (ERR)
and glutamic acid-arginine-arginine (ERRR)) were anchored to
a SNS-type carboxylic group containing monomer in order to
increase their solubility in water. This modification allows well-
All amino acids and reagents for peptide synthesis were obtained
from Chem-Impex, IL, USA. Dimethylformamide (DMF), CH
CH Cl and starting materials and reagents for the SNS monomer
synthesis were purchased from Sigma-Aldrich. Glucose oxidase
GOx, β-D-glucose: oxygen 1-oxidoreductase, EC 1.1.3.4, 17300
3
CN,
2
2
(
organized molecular structure of the conjugated polymers on units/g solid) from A. niger and glucose were purchased from
the substrates, and they constitute wonderful three Sigma-Aldrich. Glutaraldehyde (GA) was obtained from Sigma–
dimensional matrices for the biomolecule deposition while Aldrich Co., LCC. (St. Louis, USA). For enzyme immobilization, a 50
maintaining their biological activity for a long term period.
Additionally, short peptide sequence allowed easy synthesize
in small quantities, making the scale up of the production of
the device economically affordable. The design makes the
biosensor an ideal candidate to develop a cheap device.
Moreover, the corresponding monomer design was really
attractive due to their functional groups as they are open to
covalent bonding. Peptides are acting as linkers between
polymer and glucose oxidase to generate a biocompatible
mM, pH 7.0 phosphate buffer saline (PBS) solution consisting of
.025 M Na HPO and 0.025 M NaH PO (Fisher Scientific Company)
0
2
4
2
4
was used. As the substrate, glucose solution (0.1 M) was prepared
by dissolving 0.18 g of glucose in 10 mL pH 7.0 PBS solution.
Measurements
Peptide synthesis was performed by the Discover Bio - Manual
medium between the enzyme and the electrode surface. To Microwave Peptide Synthesizer (CEM Corporation, USA). The
access the conjugated biopolymer architecture, the purification of the crude peptides was performed with Dionex
constructed monomers (SNS-COOH, SNS-ERR, and SNS-ERRR) Ultimate 3000 Series equipped with a variable wavelength
were electrochemically polymerized on the graphite electrode
via cyclic voltammetry. The electropolymerization of the
monomers comprises an E(CE)n (electrochemical, chemical,
electrochemical) mechanism consisting, as the first step, the
absorbance detector using a reverse phase C8 column
Hypersil Gold, 12 μm, 250 × 10 mm). A binary gradient of
(
water (0.1% TFA) and acetonitrile (0.1% TFA) were used with a
−1
flow rate of 3.0 mL min and the eluent was monitored by UV
2
6,27
formation of a radical cation.
After electropolymerization
absorbance at 210 and 280 nm. Fractions were collected and
lyophilized after their purity was confirmed by analytical HPLC
performed using a RP-C18 column (Acclaim 120, 3.0 μm, 4.6 ×
of the monomers, GOx was immobilized on these polymeric
surfaces and used for glucose detection. In a subsequent step,
the matrix was fixed using glutaraldehyde (GA) as the
crosslinking agent. In this study, it is demonstrated that
conjugated coatings based on peptide sequences in
combination with glucose oxidase provide a simple route to
manufacture surfaces that can act as a glucose biosensor.
Scheme 1 displays the procedure to construct glucose
biosensor.
−
1
1
50 mm) with a flow rate of 0.5 mL min .
PalmSens potentiostat (PalmSens, Houten, The Netherlands)
was used for the amperometric measurements. Three
electrode system containing a graphite rod (Ringsdorff Werke
GmbH, Bonn, Germany, type RW001, 3.05 mm diameter and
1
3% porosity), a platinum wire (Metrohm, Switzerland) and a
Ag wire were used as the working, counter and reference
electrodes, respectively. Prior to electrochemical
polymerization reaction, spectroscopic grade graphite rods
were polished on an emery paper and washed thoroughly with
2
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distilled water. For biosensor applications, electrochemical The purification was done by semi-preparative RP-HPLC
polymerizations were carried out on cleaned graphite (Figures S3 and S4) and characterizations were performed by
1
electrodes via cyclic voltammetry using Gamry Instruments LC-MS and H NMR spectroscopy to confirm the structures.
Reference 600 potentiostat/galvanostat (GAMRY Instruments
HRMS for 4; calculated: 792.3074, Found: 792.3077.
Inc., Pennsylvania, USA). During electrochemical and
IR for 4 (ν): 3306, 3188, 3500-3000 (bs), 1647, 1532, 1501,
spectroelectrochemical studies, indium tin oxide (ITO) doped
−
1
1
414, 1180, 1045, 857, 699 cm .
glass slides were used. Spectroelectrochemical studies were
performed via UV-Vis spectra (Agilent G1103A
1
6
H NMR for 4 (d -DMSO): 8.92 (d, J=6.7 Hz, 1H, -NH-CH), 8.45
spectrophotometer). The potential for spectroelectrochemical (d, J=6.7 Hz, 1H, -NH-CH), 8.06 (d, J=8.2 Hz, 2H), 7.42 (d, J=8.4
studies was controlled using Solatron 1285 Hz, 2H), 7.31 (dd, J=5.0, 1.2 Hz, 2H), 7.16 (bs, 2H, -NH ), 6.97
potentiostat/galvanostat. All experiments were conducted (bs, 2H, -NH ), 6.89 (dd, J=5.3, 3.6 Hz, 2H), 6.66 (dd, J=3.6, 1.2
a
2
2
under ambient conditions (25°C). In amperometric analyses, Hz, 2H), 6.59 (s, 2H), 4.45-4.52 (m, 1H, αH), 4.05-4.10 (m, 1H,
the data were given as the average of three measurements CH), 3.90-3.95 (m, 1H, CH), 2.91-3.22 (m, 4H, -CH -N), 2.30-
2
and standard derivations were recorded as ±SD. For surface 2.36 (m, 2H, -CH -COOH), 1.80-1.85 (m, 2H, -CH -), 1.49-1.66
2
2
imaging of the fabricated biosensor, scanning electron (m, 8H, -CH -).
2
microscope (SEM) (JEOL JSM-6400 model) was used.
HRMS for 5; calculated: 947.4007, Found 947.4018.
IR for 5 (ν): 3281, 3194, 3500-2950 (bs), 1647, 1533, 1501,
−1
Synthesis of SNS-COOH and peptide anchored SNS-type 1414, 1170, 1042, 857, 695 cm .
monomers (SNS-ERR and SNS-ERRR)
Methyl 4-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl) benzoic acid (d, J=7.7 Hz, 1H, -NH-CH), 8.05 (d, J=7.5 Hz, 1H, -NH-CH), 7.99
SNS-COOH)
(d, J=8.4 Hz, 2H), 7.93 (d, J=7.7 Hz, 1H, -NH-CH),7.49 (d, J=8.4
SNS monomer was synthesized based on a literature report in Hz, 2H), 7.33 (dd, J=5.1, 1.0 Hz, 2H), 7.23 (bs, 1H, -NH), 7.13
1
H NMR for 5 (d -DMSO): 8.72 (d, J=7.2 Hz, 1H, -NH-CH), 8.22
6
(
2
8
three steps starting from thiophene and succinyl chloride.
The first step was the Friedel Craft acylation reaction between 6.68 (dd, J=3.6, 1.1 Hz, 2H), 6.60 (s, 2H), 4.49-4.47 (m, 1H, αH),
thiophene and succinyl chloride using AlCl
at 0 °C to give 4.24-4.30 (m, 2H, CH), 4.14-4.19 (m, 1H, CH), 3.04-3.16 (m, 6H,
diketone 1. The following step was the Paal-Knorr Pyrrole -CH -N), 2.34-2.42 (m, 2H, -CH -COOH), 1.64-1.74 (m, 2H, -CH
(bs, 2H, - NH ), 7.10 (bs, 1H, -NH), 6.90 (dd, J=5.1, 3.6 Hz, 2H),
2
3
-
2
2
2
synthesis with methyl-4-aminobenzoate using pTSA as catalyst ), 1.46-1.58 (m, 12H, -CH
-).
2
1
in toluene at reflux temperature under argon. As the last step,
H NMR and IR spectra of 4 and 5 are quite similar since there
1
is only one extra arginine in the conjugate 5. In H NMR
basic hydrolysis of methyl ester with LiOH in THF/MeOH
mixture at room temperature gave SNS-COOH monomer 3 in
good yield (26% over 3 steps) (Scheme 2). The structures of
spectrum of both conjugates, the peaks between 6.6 and 8.0
ppm are due to the aromatic hydrogens on the SNS moiety of
the molecule. α-Hydrogens of the peptide backbones are
between 3.9 and 4.5 ppm. Aliphatic hydrogens from glutamic
1
monomers were confirmed by H NMR, IR spectroscopy, HRMS
2
and compared with the literature results.
9
−1
IR (ν): 3430–3187 (br, OH), 1660 (C=O), 1510, 1417 cm (C–S acid and arginine residues are seen between 1.5 and 2.0 ppm
in thiophene); H NMR (CDCl
) 6.50–6.53 (m, 4H), 6.89 (dd, as multiplets. In IR spectra of both conjugate 4 and 5, there are
H), 7.16 (dd, 2H), 7.38 (d, 2H), 7.69 (d, 2H); MS (EI) m/z = 351 the characteristic C-S band around 1414 cm , amide carbonyl
1
3
-1
2
+
-1
(M+, 100%); HRMS (ES+) calcd. for C19
2 2
H13NO S
[M+1] bands around 1650 cm and NH and OH bands around 3300
-1
352.0464, found 352.0458.
cm (Figures S1 and S2).
Scheme 2. Synthesis of SNS-COOH based on the literature protocols.
SNS-Peptide Conjugate (SNS-ERR (4) and SNS-ERRR (5))
To synthesize peptide moiety, solid phase peptide synthesis
based on standard Fmoc chemistry was performed using Rink
Amide resin under microwave conditions. 0.05 mmol resin
were used and each coupling was done at 20 W and 75 °C with
HBTU in DMF as the coupling agent and DIEA as the base. 5.5
fold amino acids were used for a 5 min. coupling duration.
Deprotection of Fmoc group was done with 20% piperidine in
DMF solution (Scheme 3).
After completing all coupling reactions of amino acids, SNS-
peptide conjugation was done on resin using the same
coupling condition for the amino acids. After the cleavage of
2
side chain and resin with 95% TFA-2.5% TIPS-2.5% H O, SNS-
COOH monomer conjugated peptide derivatives (4 and 5) have
been obtained as the crude products.
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Scheme 3. Synthesis of SNS conjugated peptides 4 and 5 using standard Fmoc
based solid phase peptide synthesis using Rink Amide resin.
Results and discussion
Electrochemical and spectroelectrochemical properties of
poly(SNS-ERR) and poly(SNS-ERRR)
Electrochemical Properties
Electroactivities of monomers were investigated using cyclic
voltammetry technique. Cyclic voltammograms of pristine
SNS-COOH monomer were obtained in ACN solution using
4 4
NaClO -LiClO (0.1 M) as the supporting electrolyte due to the
poor solubility of the monomer in water. The electrochemical
oxidations of SNS-ERR and SNS-ERRR were determined in a
-
2
-1
solution of 1x10 molL monomer and 0.1 molL NaClO
-1
4
-
LiClO /H O electrolyte–solvent couple at a scan rate of 100
4
2
mV/s. Water solubility of the proposed monomer-peptide
conjugates enable several advantages. One of the main
advantages of such system is that SNS-ERR and SNS-ERRR can
be easily electropolymerized in aqueous solutions that
provides easy production of polymers via green chemistry.
Additionally, electrodeposition could be performed at
relatively low positive potentials.
Electropolymerization of SNS-COOH was performed between
0.0 V and +1.0 V in non-aqueous solution. The electroactive
film was developed on the ITO surface as accompanied with an
increase in the current. An irreversible oxidation peak was
observed at +0.72 V referring the monomer oxidation. Upon
anodic scans, polymer oxidation and reduction peaks evolved
at +0.58 V and +0.30 V (Figure 1A).
Electrochemical Biosensor Fabrication and Amperometric
Measurements
Initially graphite electrodes were coated with SNS-ERRR via
electrochemical deposition according to the same procedure
used
for
biosensor
preparation.
Electrochemical
polymerization of the monomer (SNS-ERRR) was performed in
-1
In the first cycle of voltammograms shown in Figure 1B, an
irreversible monomer oxidation peak for SNS-ERR was
centered at +0.30 V while reversible polymer redox waves for
poly(SNS-ERR) were observed at +0.25 V and +0.14 V
accompanied with an increase in the current density. Since
polymerization has been carried out potentiodynamically ecah
run exhibits a higher current density. This is due to the
electrode area increasing at every run. In Randles Sevcik
equation all the parameters are fixed whereas the electrode
area is increasing during polymer coating on the surface. Since
the polymer itself is electroactive, this may last until all
monomer is consumed. Hence the first cycle refers to the least
current whilst the last one stands for the highest current. The
reversible redox couple proves that electrochemical
polymerization is proceeding at the ITO electrode surface to
form an electroactive polymer film. An irreversible anodic
wave for SNS-ERRR at +0.32 V stands for monomer oxidation
−
3
-1
a cell containing 10 molL of monomer, 0.1 molL solution of
NaClO /LiClO (1:1) as the supporting electrolyte in the
4
4
potential range between -0.3–0.5 V. Afterwards, it was used as
the functional platform containing peptide handles on the
surface for biomolecule deposition. For the GOx
immobilization, an enzyme solution (0.75 mg in 10.0 μL, 50
mM potassium phosphate buffer, pH 7.0) was spread over the
polymer coated surface. Then, 1.0 % of 5.0 μL glutaraldehyde
in potassium phosphate buffer (50 mM, pH 7.0) were cast on
the electrode as the cross linker, and the electrode was
allowed to dry at ambient conditions for 120 min. Then,
loosely bound enzyme molecules were removed by rinsing the
electrode surface with the distilled water. It was stored at 4 °C
when not in use. All amperometric measurements were
carried out at room temperature in a reaction cell filled with
10 mL PBS, pH 7.0 under mild stirring and biosensor response
signals were recorded as the current (μA) via following the
oxygen consumption at −0.7 V due to the result of enzymaꢀc
activity in the bioactive surface. The calibration curves were
obtained by plotting the current signal values of a series of
standard glucose solution (μA) against different glucose
concentrations. After the current was reached equilibrium, a
certain amount of analyte solution was injected into the
reaction medium, then equilibrium was established again. The
difference between the two constant current values gave the
biosensor response.
(
Figure 1C). Cyclic voltammetry studies indicated that SNS-ERR
get oxidized at lower potentials than SNS-ERRR. Slight
differences in the oxidation potentials of monomers are
probably due to long chain peptide moieties in the monomer
structure. During anodic scan a doping/dedoping couple was
detected for poly(SNS-ERRR) as +0.23 V and +0.12 V at a scan
rate of 100 mV/s.
4
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Fig. 1. Repeated potential scan electropolymerization of (A) SNS-COOH in
NaClO -LiClO /ACN electrolyte–solvent couple (B) SNS-ERR and (C) SNS-ERRR in
.1 M NaClO -LiClO /H O electrolyte–solvent couple at a scan rate of 100 mV/s
up to 15 cycles).
Fig. 2. Spectroelectrochemistry of (A) poly(SNS-ERR) and (B) poly(SNS-ERRR)
polymer films on ITO coated glass electrodes in monomer-free 0.1 M NaClO
LiClO /H O at various applied potentials.
4
4
4
-
0
(
4
4
2
4
2
Table 1. The colors and color coordinates of conducting polymers in accordance with
CIE standards.
Poly(SNS-ERR)
.0 V 0.5 V
Poly(SNS-ERRR)
0.3 V
0
0.0 V
0.5 V
Spectroelectrochemistry
Spectroelectrochemical studies were carried out to determine
the optical properties of poly(SNS-ERR) and poly(SNS-ERRR)
upon applied potentials. For these measurements, polymer
films were electrochemically deposited on ITO-coated glass
L: 92.734
L: 84.602 L: 88.019
L:87.953
L:85.681
a: -10.762 a: -3.040
a: -8.261
a: -6.259 a: -1.757
b:2.716
-
2
-1
-1
plates from 1x10 molL monomer solution in 0.1 molL
NaClO -LiClO /H O. Changes in optical properties of poly(SNS-
b: 21.783 b: -1.057 b: 45.918 b:18.678
4
4
2
ERR) and poly(SNS-ERRR) were investigated in a monomer free
solution via applying increasing potentials. The newly
produced electronic bands were recorded as a function of
applied potential (Figure 2). The neutral state absorption
maxima of poly(SNS-ERR) and poly(SNS-ERRR) were observed
as wide absorptions at 470 nm and 432 nm, respectively. The
maximum absorptions in the visible region can be attributed to
the high energy π- π* transition of neutral state polymer.
Upon incremental stepping of the potential from 0.0 V to +0.5
V, the absorption of π- π* transitions decreases, and two new
optical transitions appear at lower energy, corresponding to
the polaronic and bipolaronic charge carriers. The charge
carriers led to different coloration for polymer films by
generating new energy transitions. As the potential increases,
the absorption peaks at about 537 and 915 nm for poly(SNS-
ERR) and 667 nm and 1055 nm for poly(SNS-ERRR), which
correspond to polaron and bipolaron species, respectively, also
increase.
Optimizations and Surface Characterization of the Biosensors
In order to obtain a proper orientation and effective binding of
the enzyme on the polymer surface, optimum polymer film
thickness was investigated by adjusting the cycle number
during electropolymerization. The cycle number in
polymerization determines the deposited charge on the
polymer film, which is very much related with the thickness of
the polymer layer. Stability of the enzyme on polymer film is
affected by the distance between the active site of the enzyme
molecule and transducer layer. Too thick polymer films hinder
the electron transfer between the enzyme and the electrode
3
0
causing low charge transfer rate. On the other hand, thin
polymer films cannot protect the enzyme from the
environmental effects which may cause deformation of the
matrix. To find the optimum polymer film thickness, the
monomer was polymerized on a graphite electrode with
different scan numbers (30, 40, 50 and 60 scans) by keeping all
the other parameters constant. The highest biosensor
response was obtained with 50 scans which corresponds to 39
Colorimetry Studies
2
nm (charge involved in the film formation is 25.4 mC/cm ) in
Colorimetry studies were performed to evaluate the
colors. Both polymers exhibited electrochromic properties.
Poly(SNS-ERRR) reveals multi-colored electrochromic behavior
with three distinct colors. The color analyses were done
according to CIE 1931 Yxy color space. For the polymers color
states for the pristine and oxidized forms are shown in Table 1.
thickness (Figure S5A).
Then, the enzyme amount was optimized in order to obtain
the highest biosensor response. If there is an excess loading of
enzyme, the immobilized enzyme may not be fixed properly on
the electrode surface since the matrix has an enzyme loading
capacity. This may result in leaching out from the surface. On
the other hand, if the enzyme amount is not sufficient, the
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3
1
desired sensitive responses cannot be recorded. In order to
optimize the enzyme amount, different electrodes having 0.5
mg (8.65 U), 0.75 mg (12.98 U), 1.00 mg (17.3 U) and 1.25 mg
Fig. 3. CVs of the bare graphite electrode, poly(SNS-COOH)-coated electrode, and
, at a scan rate of 100
3-/4-
poly(SNS-COOH)/GOx electrode (in 5.0 mM Fe(CN)
mV/s)
6
(21.63 U) GOx were prepared. When their responses were
compared, the highest signal and sufficient stability were
obtained with 0.75 mg GOx (Figure S5B).
Finally, the optimum value for the pH value of the buffer
solution was determined to provide an effective biosensor
platform since enzyme molecules are affected by pH
3
2
changes. Amperometric measurements were performed
using different buffer solutions in the range of pH 5.0–8.0
(sodium acetate buffer at pH 5.0, sodium phosphate buffer at
pH 6.0-7.5, tris buffer at pH 8.0, 25 °C). The highest enzyme
activity; therefore, the highest stable biosensor response was
obtained with pH 7.0 and used for further experiments.
Cyclic voltammetry (CV) was used to investigate the charge
transfer process on the biofilm surface. Effective electroactive
surface areas for each surface modification were calculated
using the Randless-Sevcik equation:
5
1/2 3/2 1/2
n v C
Ip = 2.69 × 10 AD
where n is the number of electrons involved in the redox
Fig. 4. SEM images showing the surface characteristics of (A) poly(SNS-COOH),
reaction, A is the electrode area (cm ), D is the diffusion (B) poly(SNS-ERR), (C) poly(SNS-ERRR), and (D) poly(SNS-ERRR)/GOx via under
optimized conditions.
2
2
−1
coefficient of the molecule in solution (cm s ), C is the
concentration of the probe molecule in the bulk solution (mol
−3
−1
cm ), and v is the scan rate (V s ). According to the equation,
increase in the peak currents can be attributed to an increase
in effective surface area. For this purpose, cyclic voltammetry
responses at bare electrode and modified electrodes were
3
-/4-
studied in a solution containing 5.0 mM Fe(CN)6 , 0.1 M KCl
and 50.0 mM PBS pH 7.0 at the potential between 0 and 1.0 V
-1
with a scan rate of 100 mV s . The voltammograms of
different electrodes; bare graphite, poly(SNS-ERRR), and
poly(SNS-ERRR)/GOx under optimum conditions were given in
Figure 3. The electroactive surface areas for bare graphite,
poly(SNS-ERRR) and poly(SNS-ERRR)/GOx modified electrodes
2
2
2
were calculated as 0.072 cm , 0.11 cm and 0.09 cm ,
respectively. In Figure 3, the poly(SNS-ERRR) modified graphite
electrode had a higher peak current (121.4 µA) than the one
on bare electrode (80.88 µA). This result revealed that the
increase in peak current can be attributed to the increase in
the effective surface area due to the generated conducting
polymer film on the electrode surface. By this way, the
released electrons as a result of the enzymatic reaction can be
easily transferred to the transducer surface which is important
for the biosensor response. After biomolecule immobilization,
the decrease in the peak current confirmed the proper
attachment of the biomolecule on the surface which
diminished the electron transfer properties because of a
possible diffusion layer. The results clearly confirm that the
immobilization step for biosensor fabrication was achieved
successfully.
(
Figures 4A-D). In the case of the typical cauliflower like
structure of conjugated polymer (poly(SNS-COOH)) (Figure 4A),
peptide integrated polymer surfaces (poly(SNS-ERR) and
poly(SNS-ERRR)) showed completely different surface
morphologies ((Figures 4B-C). A homogenous film observed on
the graphite surface in the case of the combination of biolayer
with the poly(SNS-ERRR) provided more ordered like structure
Scanning electron microscopy (SEM) technique was used to
assess the surface morphology of the modified electrodes
(Figure 4D). This ordered structure allows to get more reliable
6
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3
and stable biosensor responses that was proved by the fabricated by Kang et al. This biosensor construction showed
6
app
M
-
analytical results.
K
1
value of 5.35 mM and the sensitivity value of 21 µAmM
-2
cm . Another glucose biosensor based on the glucose
Analytical Characterization of the Biosensors
oxidase/one-dimensional hierarchically structured TiO
2
-
Amperometric responses of the biosensor were recorded by
adding different concentrations of glucose to buffer solution.
The characteristic calibration curves for glucose were plotted
and responses of biosensors were shown in Figure 5. Linear
ranges for each biosensor were determined as 0.1–0.5 mM for
the poly(SNS-COOH)/GOx and the poly(SNS-ERR)/GOx and
-1
modified electrode proposed the sensitivity of 9.9 µA mM cm
app
2
37
value of 1.54 mM. Poly(SNS-ERRR) matrices
and the K
M
exhibited higher affinity toward the glucose substrate; hence,
the designed biosensor served our purposes perfectly.
Furthermore, a glucose oxidase biosensor prepared with
methyl viologen-mediated, as the redox mediator, and
glutaraldehyde (GA) crosslinker in the presence of bovine
0.01-0.75 mM for poly(SNS-ERRR)/GOx modified electrodes
with the correlation coefficients of 0.984, 0.997 and 0.998,
respectively. Sensitivity values of the biosensors were also
3
8
serum albumin (BSA) as carrier protein gives LOD as 20 µM.
In an another example, based on glucose oxidase immobilized
by glutaraldehyde co-crosslinking with bovine serum albumin
−
1
−2
calculated as 19.62 μAmM
cm
for the poly(SNS-
−1
−2
COOH)/GOx, 42.0 μAmM cm for the poly(SNS-ERR)/GOx
−1
−2
and Nafion® cation-exchange polymer, on
phthalocyanine–cobalt(II) tetra(5-phenoxy-10,15,20-
triphenylporphyrin), (CoPc–(CoTPP) ) pentamer film modified
a cobalt(II)
and 91.37 μAmM cm poly(SNS-ERRR)/GOx, respectively.
Poly(SNS-ERRR) modified biosensor showed the best results
among the three biosensors as given with the equation;
4
3
9
2
glassy carbon electrode (GCE), the LOD was found as 10 µM.
y = 6.2875x + 0.0571 (R = 0.998). For optimum biosensor, the
These results also confirm the proper surface design in the
present study to achieve a very low LOD without any need
for membrane, redox mediator or carrier protein. The
designed peptide-modified conjugated polymer layer shows
both mediator and carrier property for enzymatic reactions
and immobilization, respectively. This is one of the main
advantages of the sensing platform for real-time applications.
Other advantages of the system are its easy preparation (not
requiring a multi-step preparation) and a fast response time.
When compared to other modifications summarized in Table
limit of detection (LOD) value was calculated by setting the
intercept of the linear range of the calibration curve to zero
using S/N (signal-to-noise ratio) = 3 criterions as 4.69 µM.
When compared to other modifications, this biosensor showed
also very low limit of detection (LOD) value (92.7 µM for
poly(SNS-COOH) and 72.14 µM for poly(SNS-ERR) modified
electrode). The results revealed that the poly(SNS-ERRR)
modified biosensor yields enhanced stability and affinity
towards the substrate than to that of the other modifications.
Furthermore, this modification displays that enzyme molecules
were anchored more precisely and well-oriented, and hence,
the analytical parameters were turned out to be better.
Additionally, the biosensor has a rapid and sensitive response
to glucose and reaches a steady-state equilibrium current in 2
s that allows easy detection of glucose in samples.
2, the response time of our biosensor was also shortened by
the help of surface orientation of the enzyme molecules. This
easy preparation procedure is easy to apply, simple and
effective for detection of glucose. A detailed comparison of the
properties of the biosensor is summarized in Table 2.
Fig. 5. Calibration curve for glucose (in 50 mM phosphate buffer, pH 7.0, 25 °C).
Kinetic parameters of the proposed biosensor were Error bars show standard deviation of three measurements.
investigated using Lineweaver–Burk plot (1/I vs 1/[S]) at
constant temperature and pH 7.0 and are listed in Table 2
comparatively with literature results. From the Lineweaver-
app
Burk plot, K
M
M
value was calculated as 0.208 mM. K shows
the affinity of the enzyme molecules to their substrate. When
compared to other studies in the literature, the present
biosensor showed a considerably low LOD, high sensitivity and
app
low K
M
values. Liu et al proposed a glucose biosensor based
on water-dispersible chitosan-functionalized graphene (CG)
3
3
and further modified it with Fe
3
O
4
.
The proposed biosensor
-1
-2
has a sensitivity value of 5.658 µAmM cm with a detection
limit of 16 µM. In another study, a biosensor was developed
for the detection of glucose with an electrode modification of
3
4
PdNPs-electrochemically reduced graphene oxide.
The
biosensor has the Michaelis constant of 5.44 mM. Adronov et
al designed a biosensor by entrapping glucose oxidase within
the poly[3-(3-N,N-diethylaminopropoxy)thiophene] and single-
app
walled carbon nanotubes films and they obtained the K
3
M
5
value of 3.4 mM. Moreover, a glucose biosensor utilizing
polyaniline and chitosan-coupled carbon nanotubes was
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Table 2. Comparison of the various glucose biosensors for the detection of glucose
app
M
Matrices on electrodes
Poly(SNS-ERRR)/GOx
LOD
Sensitivity
K
References
-
1
-2
(
μM)
4.69
5.9
35
(μAmM cm )
91.37
4.2
(mM)
0.208
NR
This work
GOx/Au/MXene/Nafion/GCE
Graphene–AuNPs–GOD
40
41
42
43
44
45
46
47
48
49
50
NR
4.73
1.48
0.97
1.59
NR
Au cypress/PB grid
NR
20
74.3
PA-g-PEG/GOx/graphite
47.72
NR
Poly(SNS-anchored carboxylic acid) /PAMAM-G4
GOD/graphitic-nanocage modified GCE
BNNTs-PaniPt-GOD
2.92
8
13.3
6
19.02
65.816
NR
3.4
PET/VACNTs-Al-foil/PFLA/GOx
7.035
500
10
0.193
8
3 3
(PAH/CdTe)12(PAH/PSS) (PAH/GOD)
GOx/MWCNTs/CS/GCE
GOD/Pt/OMC/Au
13
2.2
50
12
NR
NR: Not reported
Other analytical parameters such as repeatability and shelf-life of
the biosensor were also evaluated. Shelf life of the biosensor was
measured by recording the biosensor responses for 0.5 mM glucose
during 31 days. In a period of 31 days, there was only 5.0% decrease
in the biosensor response. When not in use, the biosensor was kept
at 4◦C. The device was still active with 95% efficiency after a month
Sample Application
To test the convenience of the biosensor glucose content in
several beverages were determined. For this purpose, the
samples were injected into the cell instead of the glucose
substrate without any pretreatment. Each sample (around 10
µL) was injected into 10 mL of buffer containing cell and an
automatic dilution occurs. By this way, the detected glucose
amounts were in the range of our detected linear range. The
responses of the biosensor were recorded for each sample and
glucose contents were estimated from the calibration curve.
All these experiments were achieved under the optimum
conditions. Percent relative error was calculated and the
biosensor results were compared with those given on the
labels provided by the manufacturers. The measurements
were carried out at ambient conditions (25 °C). The proposed
biosensor design is suitable for use in real time analysis to
investigate glucose content in beverages. The obtained results
are summarized in Table 3.
(
Figure 6A). Additionally, the operational performance of poly(SNS-
ERRR)/GOx was investigated for 20 successive measurements (with
a standard deviation of 0.034% and relative standard deviation of
4
.01%). The stability of the biosensor for repetitive uses in that
period is quite good. Furthermore, the investigation of the
biosensor selectivity toward possible interfering substances is a
mandatory step in the development of biosensors. Hence, some
potential interferents like ascorbic acid and urea (0.5 mM) were
injected to the reaction cell instead of glucose as the substrate. The
biosensor responses were followed and shown in Figure 6B. No
responses were recorded for these interferents at -0.7 V potential
under the optimized conditions that proves the biosensor can be
comfortably used for the glucose determinations in various
applications even within an environment with such interferents.
Fig. 6. (A) Shelf-life analysis of the biosensor during 31 days and (B) biosensor
responses to glucose and interfering substances (in 50 mM phosphate buffer, pH
7.0, 25 °C).
Table 3. Results of glucose analyses in beverages.
Glucose Content (mM)
Sample
Product
Label
poly(SNS-ERRR)/GOx
biosensor
Relative
Error
(
%)
S® Ice tea
S® Milk
0.37
0.25
0.62
0.35 ± 0.10
0.23 ± 0.01
0.59± 0.11
5.41
8.0
C® Coke
4.83
8
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1
4 C. Alemán, A. Bianco and M. Venanzi, Peptide materials:
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Two new water-soluble monomers were designed and
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1
1
1
1
1
13
characterized using H, C NMR and FTIR spectroscopy and
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monomers were done by semi-preparative RP-HPLC. Peptide-
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on graphite electrode via cyclic voltammetry and their
1
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Moreover, the surface characteristics of the electrodes were
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2
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range of 0.01−0.75 mmolL (R = 0.998), sensitivity 91.37 25 T. Trimaille, K. Mabrouk, V. Monnier, L. Charles, D. Bertin and
app
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peptides with enzymes, the biosensor showed high sensitivity
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