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ChapterV Conducting polymer thin film

modified electrochemical sensor

for uric acid and glucose

0.0 0.2 0.4 0.6 0.8 1.0

-20

0

20

40

60Glucose response

at SPEEK/PDMA/CuNFs/GCE

I/

A

E / V

Copper nanoflower

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Chapter V: Conducting polymer............................. Uric acid and Glucose

[132]

5.1. Polyelectrolyte for uric acid detection in the presence of ascorbic

acid

Simple and rapid method for uric acid (UA) determination in the presence of

ascorbic acid (AA) and other interferents has received considerable attention [1].UA is a

primary end product of purine metabolism [2]. For a healthy person, UA concentrations

in serum and urinary excretion were 0.24-0.52 mM and 1.49-4.46 mM, respectively [3].

AA is a vital vitamin of human diet, and used for prevention and treatment of some

diseases. Thus, concentration of UA and AA in biological fluids (human urine), always

changed and reflected by body metabolism [4]. Therefore, a simple sensor for their

simultaneous detection in biological fluid or human urine is important for medical

diagnosis.

For determining UA concentration, numerous methods such as in ion-exchange

column chromatography [5], enzyme analysis [6], HPCE [7], were developed. However,

cost and complexity of these methods limited their applicability. Electrochemical

methods are quick, simple and sensitive [8]. During direct determination of UA,

interference of AA is a main problem, because their almost same oxidation potential.

High concentration of AA (as in biological fluids) often interferes during UA detection.

Thus, detection method should be more selective and sensitive especially for biological

fluids. Chemically modified electrodes were reported for determination of UA in

presence of AA. For instance, GC electrode electrochemically modified with Luminol

[9], Tyron [10], conducting polymers [11], silica gel [12], methylene blue adsorbed on a

phosphorylated zirconia-silica composite [13], thiadiazole [14], MWCNT-Nafion [15],

mesoporous carbon/Nafion composite [16], poly(Evans Blue) [17] and copper [18], were

reported in the literature. But, these methods are difficult, cost intensive, less stable and

consumed more time. UA sensors were fabricated with positively or negative charged

films, self-assembled monolayer and LBL technique using poly(dimethylsiloxane) [19],

PEDOT/SDS [20], PDDA/Uricase enzyme [21], CPB [22,23], Nafion [24] and PDDA

[25]. But, these techniques are also difficult and less sensitive. Thus, there is an

expanding demand to develop simple, reliable, sensitive, and stable sensor with improved

characteristics for simultaneously sensing of AA and UA.

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[133]

Polyelectrolytes were polymers whose repeating units bear an electrolyte group.

These groups will dissociate in water, making the polymers charged and properties are

thus similar to electrolytes (salts) and polymers (high molecular weight compounds).

They sometimes called polysalts, their solutions are electrically conductive. They were

widely used for: molecular self-assembly techniques for thin film deposition of

electrically conducting polymers [26], conjugated polymers for light emitting devices

[27], nanoparticles [28], in second order nonlinear optical (NLO) devices [29] and

bimolecular sensor [21]. Cationic polyelectrolyte (PDDA) was explored for simultaneous

detection of UA and AA at GC electrode, because of its electro-conductive nature and

electrostatic interaction with charged molecules [30,31].

Thus, we are exploiting the polyelectrolyte environment for simultaneously

detection of UA and AA at GC electrode by CV and DPV. Both (UA and AA) exhibited

well separated oxidation peak potentials (360 mV). This method was also validated for

human urine samples with 3.0–8.0% measuring error. Developed approach offered an

effective tool for simultaneous determination of different biomolecules.

5.1.1. Materials

Ascorbic acid (AA) and poly (diallyldimethylammonium chloride) (PDDA) (low

molecular weight, 35 wt. % in water) were received from Sigma-Aldrich, while UA was

received from Chemika-Biochemika reagent, and used without further purification. All

other chemicals of AR grade were received from S.D. Fine Chemicals (India).

5.1.2. Electrochemical studies

All electrochemical experiments were performed using a potentiostat/galvanostat

Autolab®PGSTAT-302 N (Eco Chemie B.V., The Netherlands), in single-compartment

three-electrode cell. GC electrode (3.0 mm diameter) was polished with 0.5 µm alumina

slurry on a flat surface, rinsed thoroughly with deionized water, and sonicated for 2 min

immediately before its use and worked as working electrode . Platinum wire was used as

an auxiliary electrode. All potentials were recorded using saturated calomel reference

electrode (SCE). Aqueous solutions were prepared through deionized water (ρ > 18.2

MΩ, Millipore Milli-Q system) deoxygenated by purging with nitrogen gas prior to start

of each experiment.

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[134]

5.1.3. Analysis of urine

The utilization of the proposed method in natural sample analysis was also

investigated by direct analysis of human urine samples, collected from our lab workers.

UA was dissolved in urine to make a stock solution (50 ml) with 5 mM concentration.

Standard successive additions of 1 µl of 5 mM UA in urine were added to the buffer pH:

4.0 containing 0.028% of PDDA.

5.1.4. Results and discussion

5.1.4.1. Electrochemical oxidation of AA and UA

Cyclic voltagramms for AA oxidation at glassy carbon electrode in 50 ml of 0.2 M

phosphate buffer (pH: 4) without or with PDDA (cationic polyelectrolyte) were

investigated (Fig. 5.1.1(A)). The electrochemical oxidation of AA is dependent on the pH

of solution. In acid solution, only oxidation peak for AA was observed, whereas in basic

solution a second oxidation peak was also observed. At pH 4.0, the oxidation of AA

involved two electrons and protons (pKa of AA: 4.17) [32]. While at high pH (>4.17),

oxidation process involved loss of a proton and two electrons due to anionic nature of AA

[32,33]. Oxidation of AA at GC electrode is irreversible and requires high over-potential.

Also, no reproducible electrode response was obtained due to electrode fouling by

adsorption of AA oxidation product. AA oxidizes at 250 mV vs SCE, while it showed

-0.4 -0.2 0.0 0.2 0.4 0.6 0.8 1.0-4

-2

0

2

4

6

A

b

a

I /

A

E / V S SCE

-0.4 -0.2 0.0 0.2 0.4 0.6 0.8 1.0

-4

0

4

8

B

b

a

I /

A

E / V S SCE

Fig. 5.1.1. Cyclic voltagramms for: (A) 0.5 mM AA and (B) 0.5 mM UA in 0.2 M phosphate

buffer solution (pH: 4.0) at 50 mV/s scan rate with: (a) without PDDA; (b) with of PDDA

solution at GC electrode.

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[135]

negative shift of about 50 mV with slightly enhanced peak current in the presence of

cationic PDDA. This may be attributed to favorable electrostatic attraction between the

cationic PDDA and anionic AA, which further increased with AA concentration within

the film.

At pH: 7.34 of blood and normal tissues, UA exists as urate (anionic form; pKa:

5.75) [3]. Cyclic voltagrams for UA without or with 0.028% of PDDA (50 ml of 0.2 M

PBS, pH: 4) at GC electrodes are depicted in Fig. 5.1.1B(a,b). UA oxidation was

irreversible at GC electrode or metal

electrode, while at graphite electrode it was

quasi-reversible. Electrochemical oxidation

of UA involves two electrons and protons

process that lead to formation of unstable

diimine species, which then attacked by

water molecules in a step-wise fashion and

converted to imine-alcohol and uric acid- 4,5

diol. The uric acid-4,5 diol is unstable and

decomposes, in absence of PDDA, UA

oxidized at 500 mV vs SCE, while negative

shift of about 50 mV with small increase in peak current was observed in presence of

PDDA. At pH: 7.2, UA exists as anion, which is more hydrophobic than AA because of

its low solubility of UA in water. Thus, negative shift of oxidation potential and increase

in current response were observed due to favorable electrostatic interaction between

PDDA and anionic UA as depicted in Fig. 5.1.2 [30,31,34].

The CVs of UA at different scan rates in the presence of PDDA on GC electrodes

are presented in Fig. 5.1.3(A). Oxidation peak potential was shifted to positive direction

with and linear relationship between peak oxidation current and square root of scan rate

(10-400 mV/s) (Fig. 5.1.3(B)), suggested diffusion-controlled UA oxidation in the

presence of PDDA on GC electrode.

GCE

PDDA

UA

AA

GCE

PDDA

UA

AA

Fig. 5.1.2. Schematic drawing of poly

electrolyte electrostatic interaction between

UA and AA anion.

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[136]

The rate constant (k) and number of electrons released during UA oxidation (n) were

determined by the following equation [35]:

)1.1.5(78.301)/(log39.201pE

where k is rate constant, n the number of electrons transferred and ΔEp is the potential

peak difference. Electrochemical oxidation of UA was proceeds by a 2e-, 2H

+

mechanism, resulting unstable diimine [15-18, 22-25]. The proposed mechanism for the

electrode reaction of UA is highlighted as follows:

ΔEp values were experimentally determined and used for estimation of k values for the

UA electrochemical reaction (Eq. 5.1.2) and presented in Table 5.1.1. In these systems

strong adsorption of reactant and product are involved, the full peak width at half-height

(EFWHM) was determined by the following equation [35]:

Number of electrons transfer (n) in the electrochemical reaction are included in Table

5.1.1., which further confirmed UA oxidation in presence of PDDA, corresponds to (Eq.

5.1.2) reaction.

-0.2 0.0 0.2 0.4 0.6

-3

0

3

6

9A

a

g

I /

A

E / V ( s SCE)

8 12 16 20 241

2

3

4

5

B

I/A

Fig. 5.1.3. (A) CVs for 0.1 mM UA in 0.2 M phosphate buffer solution (pH: 4.0) in the presence

of 0.028% of PDDA at GC electrode at different scan rates: (a) 10, (b) 20, (c) 50, (d) 100, (f)

200, (e) 300, (g) 400 mV/s and (B) The relationship between the Ip and v1/2

for UA.

)3.1.5(6.90n

EFWHM

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[137]

Table 5.1.1. Variation of the parameters obtained from Fig. 5.1.3, as a function of the

potential scan rate.

5.1.4.2. Effect of PDDA concentration on AA and UA oxidation

Effect of PDDA concentration on UA and AA oxidation was explored by variations

of anodic peak potential with

current (Fig. 5.1.4). Initial change

of Epa and Ipa with addition of

PDDA, may be because of: (i)

adsorption of PDDA on electrode

surface; and/or (ii) change in

electrical activity and

concentration at electrode surface.

Adsorption of PDDA was

responsible for electrode over-

potential and electron transfer rate

[23,30,31]. After complete

coverage of the electrode surface,

PDDA formed polyelectrolyte bridging in the bulk without effecting electrode oxidation

process, which was confirmed by the unexpected change in Epa and Ipa in lower PDDA

percentage.

All experiments were carried out using PDDA solution (0.028%, w/v) (sufficient to

saturate the electrode surface with PDDA charged species). Oxidation of AA or UA at

υ/mV s-1

ΔEp/mV EFWHM/mV n k/s-1

10 67 120 0.7 0.14

20 64 12 0.7 0.30

50 66 119 0.7 0.75

100 75 119 0.7 1.30

200 121 111 0.8 1.58

300 140 103 0.9 1.93

400 163 87 1.0 1.98

0.00 0.05 0.10 0.15 0.20 0.255.1

5.4

5.7

6.0

6.3

peak potential

peak current

PDDA/%

I pa/

A

0.39

0.40

0.41

0.42

0.43

0.44

0.45

Ep

a/ V

Fig. 5.1.4. Variation in anodic peak current (Ipa) and

anodic peak potential (Epa) UA (0.5 mM) with PDDA

percentages at GC electrode, respectively.

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[138]

GC electrode is irreversible with high over-potential [36]. Thus electrode responses were

non-reproducible because of surface fouling. Due to presence of PDDA, electrode

response was stable and showed anti-fouling properties without any reduction in current

(for more than 30 cycles). In the absence of PDDA electrode surface fouling decrease

40% response after 30 cycles.

5.1.4.3. Effect of pH on AA and UA oxidation

The effect of solution pH on peak currents and potentials during AA and UA

oxidation (0.5 mM concentration each) was studied in 0.2 M phosphate buffer in presence

of PDDA (0.028%). At pH (2.0-7.0), both peak potentials shifted towards negative

direction. But, at pH (2.0-4.0), peak currents were increased, while reduced in 5.0-7.0 pH

range. This observation may be explained due to facilitated deprotonation step in

oxidation process at high pH [37]. Anodic potential for AA oxidation showed 55 mV/pH

slopes between pH (2.0-4.0) range (1e-/1H

+ oxidation) [38]. While at high pH, 30 mV/pH

slope value suggested involvement of 2e-/1H

+ [33]. Below pH: 4.17 (pKa), AA exists as

ascorbate anion (A-), thus A

- and AA may co-exist in 1:1 ratio [39]. Hence observed peak

current at pH 4.0 corresponded to the presence of the both the species in solution. 51

mV/pH slope for anodic potential in pH range: 2.0–7.0 confirmed 1e-/1H

+ UA oxidation

process [38]. UA oxidation involved two electrons to produce 4,5-dihydroxy-4,5-

dihydrouricacid (unstable) [34]. The oxidation peak current for UA was increased with

pH (2.0-4.0). Hence, pH: 4.0 in presence of PDDA is optimum condition for maximum

peak currents and peak separation (360 mV).

5.1.4.4. Simultaneous determination of AA and UA

Electrochemical sensing of biomolecules using GC electrode suffered from

interference of AA (mM concentration in biological fluids); because the fouling effect of

AA oxidation product is not negligible and the AA oxidation potential is quite close to

most of biomolecules. Herein, we tried simultaneous detection of UA and AA. In the

presence of PDDA, CV studies at GC electrode, revealed comparatively high separation

of UA and AA. The CVs for UA and AA (0.5 mM each) in 0.2 M phosphate buffer (pH:

4.0) with and without PDDA (0.028%) at GC electrodes are presented in Fig. 5.1.5. In the

absence of PDDA, both showed single peak at a potential different from their individual

oxidation potentials. At GC electrode, repetitive oxidation potential of AA was gradually

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[139]

-0.3 0.0 0.3 0.6 0.9

0.0

0.5

1.0

1.5

2.0

2.5

c

b

a

I /

A

E / V S SCE

Fig. 5.1.6. Differential pulse voltammograms for:

(a) AA (0.50 mM); (b) UA (0.50 mM); and (c)

mixture of AA and UA (0.50 mM, each) in the

presence of PDDA at GC electrode in 0.2 M

phosphate buffer solution (pH: 4.0).

-0.8 -0.4 0.0 0.4 0.8 1.2

-4

0

4

8

12

b

a

I /

A

E / V s SCE

Fig. 5.1.5. CVs for mixtures of AA and UA (0.5 mM,

each) in 0.2 M phosphate buffer solution (pH: 4) at 50

mV/s scan rate: (a) without PDDA; (b) with 0.028%

of PDDA, at GC electrode.

shifted towards anodic side, i.e.,

the non-reproducible electrode

response [36]. In absence of

PDDA, AA may affect responses

of UA at GC electrode [34]. In the

presence of cationic PDDA, clear

separation between AA and UA

peak potentials was observed. For

confirmation, the responses of UA

and AA were investigated by the

more sensitive method,

differential potential voltammetry

(DPV).

DPV curves for the oxidation of AA and UA (0.50 mM each, separately or mixture)

in the presence of PDDA in 0.20 M

phosphate buffer solution at GC

electrode are presented in Fig. 5.1.6.

Without PDDA, individual response

data for AA and UA were not

included. It was reported that in this

condition, indistinguishable oxidation

of AA and UA occurred at same peak

potential (600 mV) [22]. While in

presence of PDDA, two well defined

oxidation peaks at 100 mV and 460

mV were observed with about 360

mV separations, which were large

enough for their simultaneously determination. Observed peak separation was compared

with reported values [22,36,40-43] on different modified electrodes in Table 5.1.2.

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[140]

Table 5.1.2. Comparison of some characteristics of the different modified electrodes for

the determination of UA in the presence of AA.

Remarkable study was performed by recording DPVs curves at different UA

concentration (1.0 - 60.0 M) with 0.2 mM fixed AA concentration (Fig. 5.1.7(A)). The

peak current increased linearly with UA concentration with 0.994 correlation coefficient,

0.05 µA µM-1

sensitivity and 0.047 µA intercept. The precision (n = 6) assessed as

relative standard deviation (RSD) was 2% standard deviation (Fig. 5.1.7(B)). The peak

potential of AA was less in compare with UA, while peak currents of UA with or without

Electrode Sensitivity

(µA/ µM)

Determination

Range (µM)

Peak

Separation (V)

References

PEDOT/GC 0.48 1-20 0.365 39

OMC/Fc 0.0224 60 - 390 0.308 40

CT/CPB/GC - 2 - 600 0.260 22

Acetate buffer/GC 18.57 5 - 10 0.310 35

BBNBH /TN/ CPE - - 0.250 41

f-OMC/IL gel - 0.1 – 100 0.300 42

Polyelectrolyte/GC 0.05 1 - 60 0.360 In this work

-0.3 0.0 0.3 0.6 0.92

4

6

8A

a

g

I/A

E / V s SCE0 10 20 30 40 50 60

0.0

0.5

1.0

1.5

2.0

2.5

3.0B

I/A

UA/ M

Fig. 5.1.7. (A) Differential pulse voltammograms of AA and UA in the presence of PDDA at GC

electrode in 0.2 M phosphate buffer solution (pH: 4.0). [AA] was kept constant at 0.2 mM; [UA]

was changed from (a) 1µM, (b) 10 µM, (c) 20 µM, (d) 30 µM, (e) 40 µM, (f) 50 µM and (g) 60

µM, (B) Relationship between the anodic peak current and the concentration of UA. Linear fit

equation; Ip/ A=0.05 ([UA]/ M) + 0.067 (Ip/ A), R2 = 0.994).

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[141]

AA were almost same. Thus in presence of PDDA, oxidation of AA did not affect the UA

oxidation at GC electrode. PDDA acted as antifouling agent and covered the GC

electrode surface after formation of complex with AA oxidation product. Thus, in the

presence of AA (0.2 mM) and PDDA, sensing of UA at very low concentration level (1

µM) at GC electrode is possible. In biological conditions, AA concentration is always

higher than UA, thus about 200 times concentration ratio of AA and UA was undertaken

to mimic the physiological and biological conditions. Observations revealed the

feasibility for simultaneous detection of AA and UA in the presence of PDDA, as

separation of peak potentials were rather appreciable.

5.1.4.5. Effect of interfering substances

Table 5.1.3. Effect of interferents on DPV current response for sensing of UA (1µM) at

GC electrode in PDDA aqueous media.

aThe response of peak current for 1µM UA was 1.68µA in the absence of interferents. The

positive and negative signs are the increase and decrease in the current response, respectively.

In biological samples, ascorbic acid, glucose, urea, citrate and oxalate, etc. coexist

with UA and interfere in its selective detection. Thus, prior to the application of sensor in

Interferents Concentration of

interferents/mM

Change in current

response for 1µM UA/µAa

Glucose 0.1

0.2

0.3

-0.056

-0.059

-0.066

Urea 0.1

0.2

0.3

-0.025

-0.040

-0.050

Citrate 0.1

0.2

0.3

0.037

0.010

0.011

Oxalate 0.1

0.2

0.3

0.027

0.031

0.012

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[142]

biological samples we often need to check the effect of interferents on the response

specific analyte. Under the optimized conditions, 1.0 µM UA was sensed in the presence

of different interferents (0.1–0.3 mM) with 0.028% PDDA in 0.2 M phosphate buffer

solution, to explore the effect of interferents. No shift in oxidation potential of UA in

presence of interferents was detected. The oxidation current responses for UA in the

absence and presence of the interferents of 0.1–0.3 mM were checked and presented in

Table 5.1.3. In all the cases, change in peak current response for 1 µM UA in the

presence of different interferents was within ±0.03 µA. Thus, the UA and AA both can be

detected simultaneously in the presence of different interferents (similar to biological

systems) in polyelectrolyte media at GC electrode.

5.1.4.6. Reproducibility and stability test

The reproducibility of the proposed method was investigated by comparing the

peak current of DPV response to UA and AA at ten times added PDDA to solution at GC

electrode independently. For relative standard deviation (RSD) were 1.2% and 1.7% at

UA and AA concentration of 0.5 mM, respectively, which showed a satisfactory value.

On the other hand, the oxidation currents of UA and AA at this condition almost did not

change with potential scanning between 0.0 and 0.8 V (0.2 M PBS, pH 4.0) during 200

cycles (decreased less than 2%). The stability of this method was assessed by the PDDA

solution mixed with 0.5 mM UA and AA, stored air at room temperature for 24 h. The

results showed that there was no apparent decrease in current response to AA and UA

after almost 25 days.

5.1.4.7. Determination of UA in human urine samples

The proposed method was applied for the determination of AA and UA in human

urine. Table 5.1.4 presents the results obtained from three parallel measurements. UA

concentrations were detected about 2.61, 2.65 and 2.68 µM, which were in the range

human urine samples [44, 45]. Further, 1 µM of UA was added to the samples and it was

further detected with 3.0–8.0% measuring error. These observations validated the

suitability of developed sensor for detecting UA concentration in biological samples.

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[143]

Table 5.1.4. Results from simultaneous detection of AA and UA in urine samples by

proposed methoda

Urine samples

Detected

(µM)

Added

(µM)

Found

(µM)

Measurement errors

(%)

Sample 1 UA

AA

2.61

ND

1.00

0.50

3.58

0.49

3.0%

1.0%

Sample 2 UA

AA

2.65

ND

1.00

0.50

3.57

0.49

8.0%

1.0%

Sample 3 UA

AA

2.73

ND

1.00

0.50

3.68

0.48

5.0%

2.0%

ND= not detected .aThe results are average of three determinations

5.1.5. Conclusions for UA detection in the presence of AA in polyelectrolyte

In this chapter, a simple and reliable approach for the simultaneous detection of UA

and AA in cationic polyelectrolyte media (PDDA) at GC electrode is described. AA and

UA interacted with PDDA and facilitated their accumulation at GC surface, resulting in

the enhancement of oxidation current. Better separation of oxidation peaks of UA and

AA were achieved. It was also observed that UA and AA both can be detected

simultaneously in the presence of different interferents (similar to biological systems) in

polyelectrolyte media GC electrode. This method was also validated for the

determination UA in human urine samples with 3.0 – 8.0% measuring error. Reported

methodology is advantageous compared with chemically modified electrode because of

its simplicity and low cost. Furthermore, reported method is suitable for simultaneous

detection of AA and UA with good sensitivity and selectivity in biological fluids.

5.2. Ion conducting polymer-metal nanoparticle composite thin film for

glucose detection

Clinical conditions of diabetes mellitus are well known and understood, yet remain

growing concern as the prevalence of the disease increases worldwide at an alarming rate.

A number of life-threatening and life-impeding conditions greatly affect the diabetic

community, resulting in a much greater risk of cardiac, nervous, renal, ocular, cerebral

and peripheral vascular diseases. It was estimated that 2.8% of the world population was

affected by diabetes in 2000 [46]. Treatment has become a far more sophisticated science,

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[144]

with self-testing becoming increasingly more compact and accurate, and continuous

glucose monitoring now obtained from practical commercial sensors. Therefore

continuous monitoring of glucose level in blood is essential which demands the

development of a reliable, fast and ‘interference-free’ method. This has been one of the

challenging tasks for both analytical chemists and biomedical engineers for a long time.

Many researchers were focusing on the development of sensors based on the

electrocatalytic oxidation of glucose [47-50].

Enzymatic glucose sensors were good selective and high sensitive, although

enzymatic sensors offer reliable results [51], the most common and serious problem is the

instability of the enzyme in the sensors. Activity of glucose oxidase (GOx) can be

seriously affected by temperature, pH, oxygen, humidity, toxic chemicals and so on.

Furthermore, a complicated procedure, including functionalization, cross-linking,

adsorption, or entrapment, is required for the immobilization of the enzyme on the

electrode surfaces, and the activity of the GOx may decrease [52]. To overcome or

alleviate the drawbacks of enzymatic glucose sensors, more methods were developed to

determine glucose without using enzymes. Hence, attempts were made to develop

enzyme free sensors. Various metal electrodes such as Pt, Au, Cu, Bi, Hg, and Ag have

been proved to be highly electro-active in the glucose oxidation requiring a high over

potential. But, most of these electrodes were found to suffer the problems of low

sensitivity and poor selectivity caused by surface poisoning from the adsorbed

intermediates and the interference from chloride [53]. More recently, some of these

drawbacks were overcome by taking advantages of special physical and chemical

properties of different nanomaterials, such as nanoparticles [54-56], nanowires [57], and

nanoporous films [58,59]. When utilizing nanomaterials, it becomes possible to minimize

the influence of interferences such as AA and UA [60]. Nonetheless, low sensitivity and

costliness still limits the practical application of many nanomaterials based sensors.

For the electrochemical sensor and biosensor applications, the electrode modifying

material is expected to possess several characteristics such as good electron transduction

capability, physical or chemical environment for the bioactivity, easy accessibility

towards the analyte and large surface area. Literature reveals that all these important

characteristics cannot be inbuilt in a single material. Hence, there is always a demand for

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[145]

the development of composite materials, comprising two or more components, to achieve

adequate sensitivity and stability for the biosensors [51,61,62].

Polyaniline and its derivatives are superior polymeric materials with high

conductivity, good redox reversibility, strong attachment to the electrode surface, and

stability [62, 63]. It has been used in many practical applications, such as electrocatalysis

[64], chemo sensors [65], and biosensors [66]. Sulfonated Poly (ether ether ketone)

(SPEEK), sulfonated ionomer, has been mixed with conducting polymer to develop a

catalyst, gas separation, storage material and electrodialysis because of its enhanced

redox activity, conductivity, and mechanical behavior [67-69]. The incorporation of

SPEEK into polymer film minimizes the brittleness of the pure polymer, improves the

solubility and enhances the long-term stability of sensor.

Copper and copper oxide based materials were of great interest for the electro-

oxidation of glucose for a long time [70-73]. Recently, there is a great amount of interest

on copper and copper oxide nanomaterials for the enzyme free detection of glucose.

Cherevko et al. reported porous CuO electrode for glucose oxidation [74]. Xu et al.

developed copper nanoparticles functionalized with dimethylglyoxime modified glassy

carbon for the amperometric detection of glucose in alkaline medium [75]. Kang et al.

reported Cu nanoclusters deposition on layer of CNT prepared on glassy carbon

electrode. The obtained sensor showed very high sensitivity and good stability due to the

increase in electroactive surface area, the synergistic electrocatalytic activity combining

Cu nanoclusters with CNTs, and the three-dimensional porous structure of the Cu

nanoclusters [76]. But, no one attempt for copper nanoparticles incorporated with soluble

conducting polymer for glucose oxidation in alkaline medium, due to polymers as

electro-inactive in basic media and insolubility in solvent. Recently, over-oxidized

polypyrrole modified electrodes fabricated for glucose detection in alkaline medium with

good sensitivity and stability [77,78].

Every year authors are being reported on enzyme free glucose sensor, though they

lack the required sensitivity and/or selectivity or the method of fabrication may be

cumbersome. In order to further improve the sensitivity and selectivity, we propose an

enzyme free glucose sensor based on the SPEEK/PDMA/CuNFs/GCE. Nano-flower Cu

particles were growth by novel one step method. We used the anion-cation pair [SPEEK--

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[146]

Gluconolactone

Glucose

CuNFs coated PDMA /SPEEK

DMA and SPEEK mixture

DMASPEEK

2e-, 2H+

PDMA growth on SPEEK

APS

CuC

l2

Gluconolactone

Glucose

CuNFs coated PDMA /SPEEK

DMA and SPEEK mixture

DMASPEEK

2e-, 2H+

PDMA growth on SPEEK

APS

CuC

l2

Fig. 5.2.1. Fabrication of SPEEK/PDMA/CuNFs/GCE sensor.

DMA+]

as reducing agent for [CuCl2] reduction, and concomitantly deposits on/in the

PDMA and SPEEK as a nanosized Cu0 metal particles. Modified composite electrode

was then fabricated and electrochemically characterized to explore their catalytic activity

for glucose oxidation in alkaline medium. Our results demonstrate advantages such as

better sensitivity to glucose, fast current response, stability, and insensitivity to

interferences.

5.2.1. Materials and methods

D(+)-glucose, ascorbic acid, uric acid and poly (ether ether ketone) (PEEK) were

purchased from Fluka Chemicals. N,N’-dimethyl aniline (DMA), Copper chloride

(CuCl2), ammonium persulphate (APS), N,N’- dimethylacetamide, potassium hydroxide

(KOH) and hydrochloric acid (HCl) were of analytical grades and used as received.

Glucose solutions were prepared in 0.1 M KOH solution fresh at the time of experiment.

Double distilled water was used throughout the experiment.

5.2.2. Fabrication of SPEEK/PDMA/CuNFs/GCE sensor

The fabrication of sensor was presented in Fig. 5.2.1. Sulfonation of PEEK was

obtained by slow addition of dried powdered PEEK in conc. sulfuric acid by 10% (w/v)

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[147]

material to acid ratio at room temperature under constant stirring for 72 h. The moisture

contamination was rigorously excluded to ensure reproducibility of sulfonation level,

which was found to be 61% by 1H-NMR spectra and ion-exchange capacity studies. It

was believed that the presence of moisture contamination renders the formation of

pyrosulfonate intermediate to inter and intra-molecular sulfone crosslink. After the

reaction, polymer was precipitated with at least five-fold volume of deionized water and

shredded to obtain fine powder. This was then filtered and washed with water until the

last trace of acidity was removed [69]. Mechanism of nanocomposites formation was

reported earlier [62]. SPEEK (8.0 mg) was dissolved in 10 ml of N,N’-dimethylacetamide

and it was mixed with clear solution 4 mmol of DMA monomer under constant stirring in

nitrogen atmosphere for 2 h. 20 ml (1mM) of CuCl2 (in 1.0 M HCl) and 20 ml (0.2 M) of

ammonium persulphate (in 1.0 M HCl) was added drop wise for ternary mixture and

binary mixture, respectively. The color of the mixture turned into green indicating the

occurrence of polymerization of dimethylaniline and the colloidal solution was

evaporated into paste. The green paste was grounded well with a mortar and this

suspension was drop coated onto the surface of GCE (Glassy carbon electrode (GCE) was

polished by 1.0 and 0.06 μm alumina powder slurry on the emery sheet and chamois

leather, respectively. Then it was rinsed with Milli-Q water for 5-10 min and ultra

sonicate in nitric acid (1:1), acetone and redistilled water (10 min each), respectively) and

dried to obtain SPEEK/PDMA/CuNFs/GCE. For a comparative study,

SPEEK/PDMA/GCE and PDMA/CuNFs/GCE electrodes were also fabricated.

5.2.3. Electrochemical studies and determination of glucose

A conventional three electrode cell assembly was used for the electrochemical

measurements. SPEEK/PDMA/CuNFs/GCE and SPEEK/PDMA/GCE sensor were used

as the working electrode. Saturated calomel electrode (SCE) and platinum wire were used

as reference and counter electrodes, respectively. The electrocatalytic activity of

SPEEK/PDMA/CuNFs/ GCE was evaluated using cyclic voltammetry in 0.1 M KOH. A

small magnetic bar provided the convective transport. Amperometric measurements were

performed with SPEEK/PDMA/CuNFs/ GCE sensor at fixed potential (0.5 V s SCE) in

0.1 M KOH solution under continuous stirring for various concentrations of glucose in an

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[148]

Fig. 5.2.2. The FTIR spectra of: (A) SPEEK, (B) DMA/SPEEK, and (C) PDMA/SPEEK

electrochemical cell. After stabilization of the baseline, glucose solution was injected and

the plot of current versus time was recorded.

For determination of glucose in blood serum (BS), serum sample was collected

from normo-glycemic person. 1.0 ml of the serum was diluted to 5.0 ml in phosphate

buffer solution (pH: 7.0) (now referred as diluted blood serum (DBS)). Different samples

of blood serum (BS), were prepared as described; (b) 0.5 ml DBS + 0.10 mM glucose, (c)

1.5 ml of DBS + 0.58 mM glucose, and (d) 1.5 ml DBS + 0.30 mM glucose, were added

to 10 ml 0.1 M KOH, respectively and the amperometric response was recorded at 0.5 V.

5.2.4. Results and discussion for detection of glucose

5.2.4.1. Structural characterization

FTIR spectra for SPEEK, DMA/SPEEK and PDMA/SPEEK films were presented

in Fig. 5.2.2. FT-IR spectra of the SPEEK was found that the aromatic C-C stretching

vibration was observed at 1599-1412 cm-1

with sharp to medium intensity, while the

absorption at 3092-3006 cm-1

with medium intensity was attributed to aromatic C-H

stretching vibration in SPEEK. In plane C-H deformation bands were observed at 1290-

1000 cm-1

, and C-H out of plane bending at 955-845 cm-1

. Peaks at 2192-1788 cm-1

were

characteristic of 1,2-disubstituted and 1,2,4-trisubstituted aromatic moiety. The

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[149]

adsorption band at 1729 cm-1

was attributed to the carbonyl group stretching vibration,

and at 1024 cm-1

to O=S=O vibrations of sulfonic acid group. Few peaks with medium to

sharp intensity for –OH group of –SO3H were observed at 3644-3392 cm-1

. The

absorption bands associated with the aromatic ring in plane skeletal deformation

vibrations at 1601, 1510 and 1467 cm-1

.The band at 750 and 696 cm-1

, indicate that there

is head to tail coupling in the polymers. Absorption bands around at 1321 and 1243 cm-1

are due to C-N stretching vibrations of tertiary amines of DMA/SPEEK , but in the

PDMA/ SPEEK composite was observed in the higher frequency region shifted to 1359

and 1269 cm-1

[79]. Observed band at 2850-3000 cm-1

indicated CH3 groups of the

composites. SPEEK exhibited characteristic absorption bands at 3450 cm−1

(assigned to

OH vibration from sulfonic group interacting with molecular water), 1633 cm−1

(a

carbonyl absorption band), and 1050-1080 cm−1

(sulfur–oxygen (O=S=O) symmetric

vibration) [80].These results confirmed growth of polymerization of dimethylaniline in

SPEEK matrix caused by the ionic interactions of cationic form of DMA and SO3

(SPEEK).

Fig. 5.2.3(A) shows the XRD of SPEEK/PDMA/CuNFs .The diffraction pattern

clearly shows three major peaks at 43.50, 50.00, and 70.48 in the range of 10 to 80 ,

which can be assigned to the diffraction from the [111], [200],and [220] planes,

respectively, of the face-centered cubic lattice of Cu0 , which match well the standard

XRD data (JCPDS 004-0836) and no additional peaks of Cu or of other impurities or

oxidation of Cu2O/CuO were seen [81]. The XRD results clearly indicate that CuNFs

were present on the surfaces of PDMA/SPEEK. SEM was applied to observe the surface

morphology of the nanocomposites (Fig. 5.2.3(B-D)). SEM image of PDMA/SPEEK

shows the existence of globular porous structure (sizes in the range of 200 nm to 1 m);

while for SPEEK/PDMA/CuNFs composites showed nanostructure with high density,

shoe flower (hibiscus flower) shaped Cu nanocrystals were produced and randomly

distributed on polymer composite.

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[150]

a b

d

(111)

20 30 40 50 60 70 802(

Rel

ativ

e in

tens

ity

(200)(220)

A B

2 m

C

2 m

D

200nm

a b

d

(111)

20 30 40 50 60 70 802(

Rel

ativ

e in

tens

ity

(200)(220)

A B

2 m2 m

C

2 m2 m

D

200nm200nm

Fig. 5.2.3 (A) XRD of SPEEK/PDMA/CuNFs and SEM images of (B)

SPEEK/PDMA, (C) and (D) SPEEK/PDMA/CuNFs composite.

5.2.4.2. Electrochemical performance towards glucose

Cyclic voltammograms (CVs) of the modified electrodes were conducted in 50ml

0.1 M KOH in the absence and presence of the glucose, respectively. As shown in Fig.

5.2.4(A), no redox peaks can be observed at bare GCE; while at the SPEEK/PDMA/GCE

exhibits a pair of one electron quasi-reversible redox wave was observed in the absence

and presence of the glucose, respectively (Fig. 5.2.4(B)). This observation is quite

interesting. Because of PDMA provides adequate redox behavior to the electrode surface

and facilitates electron transfer at the electrode. Thus, an enhanced electrocatalytic

0.0 0.2 0.4 0.6 0.8

2

4

6

8 A

b

a

I/A

E / V S SCE0.0 0.2 0.4 0.6 0.8

-1

0

1

2

3

4

b

a

B

I/A

E / V( s SCE)0.0 0.2 0.4 0.6

-10

0

10

20

30

40

C

a

b

I/A

E/V S SCE Fig. 5.2.4. Cyclic voltammograms (CVs) of (A) bare GCE, (B) SPEEK/PDMA/GCE and (C)

SPEEK/PDMA/CuNFs/GCE in 50ml of 0.1 M KOH solution: (a) absence of glucose and (b)

presence of 1.0 mM glucose, respectively.

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[151]

0 25 50 75 1000

5

10

15

20

25

y = 0.225x + 0.75

R² = 0.996

I/A

/mV/S

0.0 0.2 0.4 0.6

-20

0

20

40

60

I/

A

E / V S SCE

Fig. 5.2.6. CVs of SPEEK/PDMA/CuNFs/GCE in

0.1M KOH (50 ml) and 0.5mMglucose at various

scan rates: 10-100 mV/s. Inset: Peak current vs. scan

rates.

current was witnessed at SPEEK/PDMA/GCE. But in above both cases of electrodes no

changes in current during the addition of glucose were seen.

In the case of SPEEK/PDMA/CuNFs/GCE, a significantly high peak current with

irreversible process than GCE and SPEEK/PDMA/GCE was observed. Thus, when

CuNFs were grafted onto

PDMA/SPEEK the electroactivity

of the electrode was significantly

improved for glucose detection

(Fig. 5.2.4(C)). PDMA and CuNFs,

synergistically contributes to the

electrocatalytic process. The above

results demonstrate that

modification of GCE by

SPEEK/PDMA/CuNFs provides

adequate electron transfer path and

enhance catalytic current for

glucose oxidation [82]. The performance of SPEEK/PDMA/CuNFs /GCE for glucose

oxidation was assessed in acidic, alkaline and neutral medium (Fig. 5.2.5). No current

response was detected in acidic or neutral medium. In alkaline medium (0.1 M KOH)

current response increased with

decrease in background noise due

relatively easy oxidation of glucose

[76]. Earlier reports also supports

our observations for enhanced

electro-catalytic activity transition

metals towards oxidation of

carbohydrates, in alkaline medium

[75,76,83] Thus, SPEEK/PDMA/

CuNFs/GCE sensor was suitable for

glucose oxidation in alkaline

medium(0.1 M KOH).

0.0 0.2 0.4 0.6

0

4

8

12

a

bcI/

A

E / V( s SCE)

Fig. 5.2.5. CVs of SPEEK/PDMA/CuNFs/GCE for

1mM glucose oxidation in different media a) 0.1 M

HCl, b) 0.1 M PBS (pH: 7) and (c) 0.1 M KOH .

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[152]

0 2000 4000 6000 80000

6000

12000

18000

24000

c

b

a

Z''/

Ohm

Z'/OhmFig. 5.2.7. Nyquist plots of EIS in 0.1 M KOH at

(a) bare GCE, (b) PDMA/SPEEK/GCE and (c)

SPEEK/PDMA/ CuNFs/GCE.

0 5 10 15 200

5

10

15

20

25

30

I/A

Cu2+

(mM)

Fig. 5.2.8. Variation of copper concentration with

peak current of glucose oxidation (0.1 M KOH,

1.0 mM glucose) at SPEEK/PDMA/CuNFs/GCE.

CVs were recorded in a solution of 0.5 mM glucose at SPEEK/PDMA/CuNFs/GCE

for different scan rates (Fig. 5.2.6). It can be seen that the potentials and peak currents of

the glucose irreversible process (anodic peak) are dependent on the scan rates. Larger

peak-to-peak separations were observed with increase in scan rates. The anodic peak

current values increased linearly with the scan rates (correlation coefficient values of

0.996 for anodic peak) in the scan rate range of 10-100 mV/s (Inset Fig. 5.2.6). The

linearity between peak current and scan rate suggests that the electrochemical kinetics is

controlled by the adsorption of glucose.

An impedance spectrum (Nyquist plot) consists of a semicircle portion observed

at higher frequency range

corresponding to the electron-

transfer-limited process and a linear

segment at lower frequencies

representing the diffusion limited

process. The diameter of the

semicircle in the Nyquist plot equals

the electron-transfer resistance (Rct),

which is related to the electron-

transfer kinetics of the redox probe at

the electrode surface. EIS of GCE,

SPEEK/PDMA/GCE, and SPEEK/

PDMA/CuNFs/GCE in 0.1 M KOH

was presented in Fig. 5.2.7. It shows

a decrease in diameter of the

semicircular portion for

SPEEK/PDMA/CuNFs/GCE and

hence the electron transfer resistance

decreases in this electrode. From the

Fig. 5.2.7, it is obvious that the bare

GCE exhibits maximum electron

transfer resistance and it decreased

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[153]

for ternary composite [84]. Variation of peak current of glucose oxidation (0.1 M KOH,

1.0 mM glucose) at SPEEK/PDMA/CuNFs/GCE with copper concentration in presence

of dimethylaniline (fixed concentration) was also investigated (Fig. 5.2.8). Further, with

increase in copper chloride concentration, peak current increased and beyond 10 mM it

limited, because agglomeration of copper particles in the polymer matrix and thus

reduction of its surface area.

As said by the related studies on the electrochemical oxidation of glucose in

alkaline media at metals modified electrodes, clear mechanism has not been proposed yet

[74,84-86].According to them, the oxidation was triggered by the deprotonation of

glucose and isomerization to its enediol form followed by adsorption onto the electrode

surface and oxidation by different oxidation states of metals. During the glucose

oxidation occurs at 0.5 V, the oxidation might not simply produce gluconic or glucuronic

acid but might instead entail C-C bond cleavage generating lower-molecular-weight

products such as formats and carbonates, which involves a 12 electron transfer. Thus, the

current response for a definite concentration of glucose has to be very high on CuNFs

modified enzyme free electrodes as compared to other enzyme free electrodes.

5.2.4.3. Amperometric sensing of glucose

To find the optimal conditions for the amperometric sensing of glucose, the effect

of applied potential on the response current of the sensor was investigated. The

amperometric curves of modified electrode in 0.1 M KOH solution with successive

addition of 0.5 mM glucose at applied potentials of 0.3 - 0.6 V was shown in Fig. 5.2.9.

In the range of 0.3 - 0.6 V the oxidation current of glucose increased with increasing

potential. With increasing potential further the side reaction takes place on the electrode,

which results in the increase of the background current and leads to the inhibition of

glucose oxidation. Thus, a constant potential of 0.5 V was chosen for further

amperometric investigations.

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[154]

100 200 300 400

0

5

10

15

20

25

30

I/A

Time/sec

0.3 V

100 200 300 400

0

10

20

30

40

Time/sec

0.4 V

100 200 300 4000

15

30

45

60

I/

A

Time/sec

0.5 V

100 200 300 400

0

10

20

30

40

Time/sec

0.6 V

Fig. 5.2.9. Amperometric response of SPEEK/PDMA/CuNFs/GCE with successive addition of

0.5 mM glucose in 0.1 M KOH (50 ml) at different applied potentials.

Amperometric measurements were performed to analyze the performance of

SPEEK/PDMA/ CuNFs/GCE sensor towards increasing glucose concentrations in 0.1 M

KOH (50 ml). Fig. 5.2.10 shows well-defined, stable, and fast amperometric responses

were observed at an applied potential of 0.5 V with successive step additions of glucose

in stirred condition (300 rpm). After successive glucose addition in step-wise, current

response increased at regular interval. It is clear that the response current of the modified

electrode increases to steady-state values upon the addition of glucose. This sensor

showed a very fast response, as the response reaches 95% of the steady-state value within

6s.

Inset of Fig. 5.2.10 shows the calibration curves for nanoflowers electrode.

Amperometric sensitivity and detection limit of the nanoflowers sensor was 19.3 µA/mM

(R2 = 0.994) and 0.1 µM with the linear range of 0.5-5 mM, respectively. CuNFs sensor

sensitivity is superior than that of the literatures for mesoporous platinum (10 µA/mM),

Cu-nanocluster/ CNT-Nafion composite (17.76 µA/mM),over oxidized polypyrrole

nanofiber/cobalt(II) phthalocyanine tetrasulfonate (5.695 µA/mM), oil/MWCNT/Cu-

nanoparticle composite (59.2 nA/mM), NiO/MWCNT (1.79 µA/mM), platinum

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[155]

nanotubes (1 µA/mM), NiO (50

nA/mM), MWCNTs (5 µA/mM),

porous Au electrode (11.8 µA/mM),

Pt-Pb alloy nanoparticles/MWCNTs

(17.8 µA/mM) and FeOOH nanowires

(12.1 µA/mM) [60,76,78, 87-94]. The

above elctrocatalytic studies of

SPEEK/ PDMA/CuNFs/GCE sensor

reveal the properties of high

sensitivity, low detection limit and

fast response time. This is attributed

to the synergetic effects by combining

Cu-nanoflowers with PDMA in the

presence of SPEEK, which greatly

increase the electrocatalytic active areas and promote electron transfer in the oxidation of

glucose.

The electrocatalytic behavior of as-prepared PDMA/CuNFs/GCE was also

investigated compared with SPEEK/PDMA/CuNFs/GCE. As shown in Fig. 5.2.11(A), an

obvious increase of anodic current of PDMA/CuNFs at 0.0 - 0.6 V can be observed when

0.5 mM of glucose was added, but the catalytic current is significantly lower than that of

0 1 2 3 4 50

25

50

75

100R

2= 0.994

Cu

rre

nt

(A

)

Glucose(mM)

0 100 200 300 400 500 6000

20

40

60

80

100b

a

0.5 mM

3.0 mM

I/A

Time (sec) Fig. 5.2.10. Amperometric responses of (a)

SPEEK/PDMA/GCE and (b) SPEEK/PDMA/

CuNFs/GCE to successive addition of 0.5 mM

glucose in 50ml of 0.1 M KOH solution at an

applied potential of 0.5 V ( s. SCE); Inset:

calibration plot of the glucose concentration vs.

current response.

0.1 0.2 0.3 0.4 0.5 0.6-20

0

20

40

A

b

a

I/A

E/V S SCE 100 200 300 400

0

20

40

60

a

b

B

I/A

Time (sec)

Fig. 5.2.11. A) CVs and (B) Amperometric curves of different electrodes for 1 mM glucose

oxidation in 0.1 M KOH, respectively; (a) PDMA/CuNFs/GCE, and (b) SPEEK/PDMA/CuNFs/

GCE.

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[156]

SPEEK/PDMA/CuNFs/GCE. The amperometric response of PDMA/CuNFs modified

electrode to successive addition of glucose is also studied (Fig. 5.2.11(B)). Compared

with SPEEK/PDMA/CuNFs/GCE, the response current of PDMA/CuNFs modified

electrode is unstable, vulnerable to external disturbances, and poorly reproducible.

5.2.4.4. Selectivity and Stability

Enzyme free sensors based on metals or alloys could easily lose their electroactivity

in the presence of Cl-, due to the formation of CuCl and/or CuCl complexes [74].

However, we did not observe any loss in electroactivity with our modified electrode

while sensing glucose in 0.1 M KOH with 0.1 M KCl; the current response remained

almost unchanged, confirming excellent stability of our sensor against chloride

poisoning.

Enzyme free sensor usually suffers from the inferences such as AA and UA,

because of these easily oxidized, have

higher electron transfer rates than

glucose. Normal physiological level of

glucose is 3-8 mM and that of AA and

UA is about 0.1 mM [61]. Therefore,

the selectivity of the modified

electrode was investigated against AA

and UA. Fig. 5.2.12 present

amperometric curve for modified

electrode in 0.1M KOH with

successive additions of AA and UA

(0.1 mM), and 1 mM glucose. As

shown in Fig. 5.2.12, the

SPEEK/PDMA/CuNFs/GCE successfully detected the glucose in the presence of UA and

AA, demonstrating almost negligible interferences from UA and AA. Similarly, there

was no obvious interference in the measure of the SPEEK/PDMA/CuNFs/GCE.

150 200 250 300 350 400

10

15

20

25

30

351 mM glucose + 0.1 mM UA

1 mM glucose + 0.1 mM AA

1mM glucose

I/

A

Time/secFig. 5.2.12. Amperometric responses of

SPEEK/PDMA/CuNFs/GCE to glucose (1.0 mM),

ascorbic acid (0.1 mM), and uric acid (0.1 mM) in

50 ml of 0.1 M KOH.

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[157]

Reproducibility and stability of the

modified electrode was investigated in

continuous operation mode. The electrode

was kept at 25 ºC in distilled water, and

tested with 0.5 mM glucose in 0.1 M

KOH for different days (Fig. 5.2.13).

Loss in current signal after first 5 days

was negligible (3%), but after 20 days,

loss in current signal was about 87.5 %.

This study has demonstrated long-term

stability and reproducibility of developed sensor, because of existing of SPEEK, has high

adhesion of CuNFs on PDMA and tightly immobilized on GCE surface.

Suitability and reliability of reported biosensor was tested in BS samples

(Preparation of BS samples described in section 5.2.3). Amperometric responses of

SPEEK/PDMA/CuNFs/ GCE to glucose solution and blood serum samples are presented

in Fig. 5.2.14 (a: 0.5 mM glucose; (b-d) blood serum in 0.1 M KOH (10 ml),

respectively). Two additions of known concentration of glucose were made and

correlated with the response obtained

in BS. At point ‘a’ addition of 0.1 ml,

0.5 mM (9.0 mg/dl) glucose solution

increases the effective concentration

with current by 9.04 µA. These

values when compared with the

values for glucose, yield

concentrations 3.51 mM (63.18

mg/dl), 4.09 mM (73.62 mg/dl) and

3.81 mM (68.96 mg/dl), their

corresponding current values were

65.1, 75.5, and 68.8 µA, respectively.

Results thus obtained were in

agreement with those from the photometric method, variation being less than 2.3%. Thus,

0 100 200 3000

50

100

150

200

250d

c

b

a

I/A

Time/sec

Fig. 5.2.14. Amperometric response of the modified

electrode for glucose solution and serum samples

tested in 0.1MKOH (10 ml) at 0.5 V applied

potential: a) 0.5 mM glucose solutions, b-d) blood

serum containing glucose concentrations.

0 5 10 15 20 25 30

88

92

96

100

I/I 0

X1

00

t (days)Fig. 5.2.13. Long-term stability of the

SPEEK/PDMA/CuNFs/GCE sensor.

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[158]

the SPEEK/PDMA/CuNFs/GCE biosensor can be used for the glucose detection in

human blood serum.

5.2.5. Conclusions for ion conducting polymer-metal nanoparticle composite thin

film for glucose detection

In the course of the current study we found that the stabilized new nanocomposites

can be formulated for enzyme free glucose sensor using a poly (dimethylaniline), SPEEK

and copper nanoparticles. The SPEEK/PDMA/CuNFs/GCE fabricated sensor

electrochemically boasts a number of advantages such as high sensitivity, good stability,

reproducibility, and quick response. Excellent selectivity was rendered by the

nanocomposites film and signals from oxidation of common interfering species such as

UA and AA can be effectively suppressed. The remarkable performance was attributed to

synergistic electrocatalytic activity combining a CuNFs and PDMA in the presence of

SPEEK, and the large electro-active surface area of CuNFs. Also, the developed sensor is

a potential candidate for routine glucose analysis in real blood serum samples. This

methodology would be extended for the development of other biosensors.

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