Comparative study of enzymatic and non-enzymatic detection of glucose using manganese ferrite nanoparticles

The use of metal oxide nanoparticles for the development of cost-effective glucose biosensors has been receiving increased attention. Enzymatic and non-enzymatic glucose sensor using polyethylene glycol (PEG) grafted manganese ferrite (PEG-MnFe2O4) nanoparticles (NPs) modified onto a glassy carbon electrode (GCE) is reported in the present study. XRD and Raman studies confirmed the cubic spinel structure of MnFe2O4. The immobilization of glucose oxidase (GOx) on PEG-MnFe2O4 (GOx@PEG-MnFe2O4) was validated using FTIR and TGA. Sensing electrodes exhibited well-defined redox peaks in 0.1 M phosphate buffered saline (PBS) solution at pH 7.4 against the reference electrode Ag/AgCl. GOx@PEG-MnFe2O4/GCE displayed a sensitivity of 1.985 μA mM−1 cm−2 in the linear range of 1 to 20 mM with a limit of detection (LOD) of 0.132 mM whereas non-enzymatic sensor exhibited a sensitivity of 1.044 μA mM−1 cm−2 in the linear range of 1 to 10 mM with a LOD of 0.099 mM. The lower Michaelis constant ( Kmapp ) value indicates greater affinity towards glucose for the enzymatic sensor. GOx@PEG-MnFe2O4 revealed selectivity specifically for glucose over various interferants such as fructose, lactic acid, sucrose, uric acid and ascorbic acid. In addition, this enzymatic sensor demonstrated better reproducibility and lifetime.


Introduction
Diabetes mellitus commonly referred to as diabetes is one of the major health concerns affecting vast majority of the world population. As estimated by the International Diabetes Federation (IDF), one in every 10 people suffers from diabetes and approximately 463 million adults in the age group of 20 to 79 years are having diabetes [1]. Diabetes is a condition resulting from the lack of insulin in body which causes abnormally high bloodglucose concentration (hyperglycemia). Monitoring the glucose level is a critical factor for the treatment of diabetes as higher dosage of medicine can trigger glucose level to drop below the normal (hypoglycemia). Reusable type glucose sensors are not commercially available and the development of such reusable sensors can revolutionize the field. Many glucose sensors (both enzymatic and non-enzymatic) are being developed in which the enzymatic type makes use of an enzyme for direct reaction with glucose aiding for electron transfer while non-enzymatic sensors cause the direct oxidation of glucose.
Enzymatic electrochemical glucose sensors have been receiving immense attention for blood glucose detection due to its high sensitivity, selectivity and low limit of detection (LOD). Among the widely used enzymes, glucose oxidase (GOx) catalyses glucose oxidation in presence of oxygen to D-glucono-1,5-lactone which then hydrolyses to gluconic acid and hydrogen peroxide whereas the enzyme glucose dehydrogenase catalyses glucose to D-glucono-1,5-lactone [2]. GOx is considered as the gold standard for enzymatic glucose sensing owing to its high specificity to glucose. GOx is a homodimeric enzyme with flavin adenine dinucleotide (FAD) bound non-covalently to its active sites [2]. However, direct electron transfer between enzymes and electrode can lead to very less sensitivity. The activity of electrode can be enhanced by conjugating nanomaterials based on carbon [3][4][5][6][7], noble metals [8] along with their alloys [9][10][11], transition metals [12] and their oxides [13,14] or alloys [15] which will aid in the electron transfer as a mediator to the sensing electrode. In spite of the fact that enzymatic glucose sensors possess high selectivity and stability, their use is restricted due to poor enzymatic activity influenced by pH, humidity, thermal conditions and presence of chemicals which degrade the enzyme [8]. These drawbacks propelled extensive research in the field of non-enzymatic sensors, out of which transition metals and their alloys have been proven to be most effective with high selectivity and sensitivity [14], however, the associated high cost limits their usage. Lee et al reported the performance of enzymatic and nonenzymatic glucose sensors using nanostructured Au-Ni alloy. This work demonstrated the superior performance of the enzymatic glucose sensor with sensitivity of 1.302 μA mM −1 with LOD of 0.29 μM validating excellent selectivity, stability and linear range compared with non-enzymatic sensor with sensitivity 0.9601 μA mM −1 with LOD of 5.84 μM [16]. In a similar study, Mohapatra et al. investigated enzymatic and non-enzymatic glucose sensors using a carbon nano-onion modified sensor possessing a higher sensitivity for the enzymatic one with 26.5 μA mM −1 cm −2 with LOD of 0.21 mM compared to the non-enzymatic with 21.6 μA mM −1 cm −2 with LOD of 0.09 mM [3]. All these experimental investigations suggested the potential of NPs-based glucose bio-sensors for improving the sensing performance with high stability, sensitivity, selectivity and LOD.
Spinel ferrites have been reported to increase the electronic conductivity, structural stability and reversibility of the electrode material which can improve the performance of electrochemical sensing devices [17]. Manganese ferrite (MnFe 2 O 4 ), one of the spinel ferrites, possesses beneficial properties such as enhanced electrical and magnetic properties with thermal and chemical stabilities, has been applied in various fields such as batteries [18], ferrofluids [19], catalysts [20] and biomedical applications [21,22]. In the present study, MnFe 2 O 4 has been chosen for glucose sensing owing to its excellent biocompatibility apart from the abovementioned advantages. In order to improve the monodispersity and structural stability of the nanoparticles (NPs), conducting polymers such as polyethylene glycol (PEG), polyethylenimine (PEI), polyvinyl alcohol (PVA), polyaniline (PANI) are being extensively investigated for biosensing applications [23]. Polymers can act as coatings which provide electrostatic, steric, or electrosteric repulsive forces between magnetic nanoparticles (MNPs) preventing aggregation and promoting NPs dispersity as well as colloidal stability [24]. Monodispersity improves the electrical conductivity of NPs which is beneficial for sensing applications. Moreover, such polymers have been utilized for the construction of biosensors as well as supporting matrix for the electrochemical activity. PEG is one of the most explored polymers as stabilizing or coating agent for NPs [25][26][27]. This hydrophilic biocompatible conducting polymer has been approved by the Food and Drug Administration for various biomedical [28] and sensing applications [29,30]. Several methods such as coprecipitation, hydrothermal, microemulsion, thermal decomposition, microwave-assisted etc have been introduced to synthesize MNPs [31,32] Among the synthesis processes, hydrothermal method possesses the advantage of the formation of monodispersed, morphologically controlled and better crystalline natured MNPs [33][34][35].
In the present study, PEG grafted MnFe 2 O 4 NPs via hydrothermal approach have been synthesized and tested towards glucose sensing. A comparative analysis for enzymatic and non-enzymatic glucose sensing has been reported. Here, PEG-MnFe 2 O 4 NPs act as mediators for promoting electron transfer in enzymatic glucose sensor whereas direct oxidation reaction of glucose results for electron transfer in non-enzymatic glucose sensor. Finally, the results of two electrodes were compared to evaluate the better sensing performance.

Materials
Glucose oxidase (GOx), D+glucose, manganese chloride tetrahydrate (MnCl 2 .4H 2 O), iron chloride hexahydrate (FeCl 3 .6H 2 O), ethylene glycol, hydrazine hydrate, polyethylene glycol (PEG)-4000, uric acid, Lascorbic acid and nafion were purchased from Sigma-Aldrich. Sodium hydroxide (NaOH), ethanol (C 2 H 5 OH), potassium chloride (KCl), potassium ferricyanide (K 3 [Fe(CN) 6 ]), sucrose, D-fructose and lactic acid were purchased from SDFCL. 10X PBS (7.4 pH) was purchased from SRL. Deionised (DI) water was used throughout the experiments 2.2. Synthesis of MnFe 2 O 4 NPs via hydrothermal method MnCl 2 .4H 2 O and FeCl 3 .6H 2 O in 1:2 ratios dissolved in 50 ml of ethylene glycol were taken in a 100 ml teflon container. 1 g PEG dissolved in 5 ml ethylene glycol was then added to the above mixture. Upon complete dissolution, 5 ml of hydrazine hydrate was added to the above mixture and maintained the pH at 12. The whole solution was stirred for 1 h under nitrogen blanket and then transferred to a stainless-steel autoclave and kept in a furnace at 200°C for 24 h. Later, the synthesized material was washed with water and ethanol several times and separated using magnetic separation. The synthesized sample is labelled as PEG-MnFe 2 O 4 .

Electrochemical analysis
Electrochemical experiments were performed using a CHI660C electrochemical analyzer with 20 ml of 0.5 M KCl+2 mM K 3 [Fe(CN) 6 ] as supporting electrolyte. The cyclic voltammetry (CV) was carried out in the potential range from −0.8 to 0.8 V for the bare GCE, PEG-MnFe 2 O 4 /GCE and GOx@ PEG-MnFe 2 O 4 /GCE at different scan rates of 10, 50 and 100 m V s −1 . CV for different glucose concentration ranging from 0-10 mM were carried out in an electrolyte of 0.1 M PBS at 7.4 pH. Differential pulse voltammetry (DPV) was also conducted at 10 mV increment for both the electrodes. Amperometric analysis was performed with the addition of glucose in an electrolyte of 0.1 M NaOH at −0.7057 V to −0.3730 V. The selectivity was further studied for enzymatic sensor by adding interferants such as fructose, lactic acid, sucrose, uric acid and ascorbic acid with the concentration of 10 mM. A schematic illustration of PEG-MnFe 2 O 4 NPs modified GCE for electrochemical sensing is shown in figure 1.

Material characterizations
Phase and crystal structure of the NPs were characterised using x-ray Diffraction (XRD) recorded in Bruker D8 Advance equipment at λ=1.54 Å from Cu K α radiation. Raman spectra providing information about the chemical bonding were recorded using a HORIBA Scientific system equipped with 532 nm laser source at 3.6 mW laser power. Surface functionalization and GOx loading was confirmed using Fourier transform Infrared (FTIR, IR Affinity-1 Spectrophotometer) spectroscopy. Thermogravimetric analysis (TGA) was performed for determining the thermal stability with the aid of a TGA, SDT Q600, TA Instruments by heating the sample from room temperature to 800°C under nitrogen environment. Morphology was studied using FEI, Tecnai G2 F30 Field Emission Gun-Transmission Electron Microscope 300 kV (HR-TEM) and a FEI, Quanta 200 Field Emission Scanning Electron Microscope (FESEM). BET and BJH isotherms were used to investigate the pore size, pore volume distribution and specific surface area employing Quantachrome Nova Station 1000 instrument. CV, DPV and amperometric analysis of the NPs were measured in a three-electrode based CHI660C

Raman spectroscopy
Raman spectrum of PEG-MnFe 2 O 4 NPs in the frequency range of 100-800 cm −1 is shown in figure 2(b). Broad peaks observed at ∼224, 262, 356 and 607 cm −1 closely match to those reported value for MnFe 2 O 4 [36]. A slight shift in the peaks towards lower wavelength can be observed which is ascribed to the longer chemical bond length of the molecule [37]. High frequency peak at ∼607 cm −1 belongs to the A 1g (Mn 2+ O) vibrational mode associated with the symmetric stretching of oxygen atoms along Mn-O bond at the tetrahedral site. Peaks observed at ∼224, 262 and 356 cm −1 correspond to the A 1g , E g and T 1 modes respectively of Fe 3+ -O bond at the octahedral site [38]. No other impurity modes are observed revealing the pristine MnFe 2 O 4 NPs. figure 3(a) displays two characteristic metal-oxygen bands at ∼872 and 545 cm −1 which can be assigned to Mn-O and Fe-O bonds at tetrahedral and octahedral sites respectively. Two bands observed at ∼1420 cm −1 and ∼1633 cm −1 correspond to the C-H bending and O-H stretching vibrations of PEG respectively [39]. A weak band at ∼2950 cm −1 corresponds to the asymmetric CH 2 bending vibration whereas a broad vibration band near ∼3346 cm −1 attributes to OH stretching vibrations of water molecules adsorbed on the surface of NPs [40]. These observed bands revealed the successful coating of PEG onto the surface of MnFe 2 O 4 NPs. FTIR spectrum (in figure 3(b)) of bare GOx exhibits a broad absorption band at ∼3280 cm −1 corresponding to the N-H stretching and peaks observed at ∼1639 cm −1 and ∼1531 cm −1 correspond to the amide bands [41]. Specifically, band at ∼1639cm −1 is observed due to the carbonyl (C=O) vibrations of peptide bonds whereas ∼1531 cm −1 is due to the N-H in-plane bending and C-N stretching modes of polypeptide chains of bare GOx.      . 73-1964). Moreover, from the FESEM micrographs (figure S1 is available online at stacks.iop.org/MRX/7/094001/mmedia), it has been confirmed that the individual NPs aggregate and form an interconnected structure resembling directional growth probably due to polymer coating and/or presence of ions like Fe 3+ and Mn 2+ . According to the compacted morphology, it is expected that the immobilisation of GOx onto the surface of PEG-MnFe 2 O 4 NPs is via physical adsorption rather than via pores.

BET analysis
The BET analysis (figure 6) suggests that PEG-MnFe 2 O 4 NPs exhibit an average pore diameter of ∼4.54 nm with a pore volume of 0.059 cc/g and a BET surface area of 52.71 m 2 g −1 . Though the exact dimensions of glucose oxidase are still uncertain, values for the dimeric structure fall within the mesoporous range (7×5.5×8 nm for the dimer) [43]. The immobilization of GOx occurs via pores when the size of the enzyme is comparable to the surface pores outside of NPs. Since the average pore size of the NPs is small compared to the dimensions of the enzyme, the immobilization of the GOx enzyme onto the surface of PEG-MnFe 2 O 4 is assumed to be accrued not via pores rather than by physical adsorption through hydrogen bonding. This hypothesis is consistent with the reported literature [44].    6 ] electrolyte in presence and absence of glucose at a scan rate 100 mV s −1 . CV curve displays oxidation-reduction peaks for all the three systems confirming the occurrence of redox reaction due to the presence of Fe 3+ /Fe 2+ in the ferricyanide solution. CV for GOx@PEG-MnFe 2 O 4 /GCE electrode exhibits the maximum current variation among others. GOx reduces flavin adenine dinucleotide (FAD) to FADH 2 which produces electrons and PEG-MnFe 2 O 4 NPs promote the electron transfer as a mediator to the sensing electrode. However, the current values are lesser for PEG-MnFe 2 O 4 /GCE electrode due to the absence of a highly reactive material (GOx) causing for electron transfer. Figure 8 shows the I-V curves of PEG-MnFe 2 O 4 /GCE and GOx@PEG-MnFe 2 O 4 /GCE with different scan rates of 10, 50 and 100 mV s −1 for a glucose concentration of 10 mM. Increasing scan rate increases the anodic current (highest oxidation current) for both enzymatic and non-enzymatic sensor. Among these, higher oxidation current of 4.097 μA was observed for GOx@PEG-MnFe 2 O 4 /GCE than PEG-MnFe 2 O 4 /GCE (oxidation current is 3.81 μA) for the highest scan rate of 100 mV s −1 . This suggests that GOx@PEG-MnFe 2 O 4 /GCE shows better electrocatalytic activity towards glucose oxidation. It is evident that the anodic peak current increases linearly with square root of scan rate with high correlation coefficient R 2 (shown in inset of figures 8(a), (b)) indicating that the reaction is diffusion controlled according to Randles-Sevcik model [4,45,46]. In addition, the electron transfer process exhibited a quasi-reversible nature. The electro-active surface area for the GOx@PEG-MnFe 2 O 4 /GCE was determined using Randles-Sevcik equation:   For determining the effective performance of the electrodes in physiological conditions, current response at varying glucose concentration of 0 to 10 mM and at a scan rate of 100 mV s −1 using 0.1 M PBS as electrolyte was investigated. It is observed that both the electrodes exhibited electrocatalytic activity towards glucose oxidation in PBS solution. Figure 9(a) depicts the CV of PEG-MnFe 2 O 4 /GCE in which the anodic peak obtained at +0.1 V and the cathodic peak obtained at −0.1 V can be ascribed to the oxidation and reduction of glucose occurring due to the direct electron transfer from the NPs to the glucose in the solution. Similarly, anodic current increase with glucose concentration evident from figure 9(b) confirms that the glucose is catalysed by the enzyme GOx using oxygen to generate hydrogen peroxide [47]. The possible mechanism of glucose sensing is supported by the following redox reactions: MnFe O 2 2 2 4 Oxidation and reduction peaks follow the reactions shown in equations (1) and (2) respectively. The overall redox reaction is presented in equation (3). In the absence of glucose, oxidation-reduction occurs due to the reversible reaction of GOx-FAD to GOx-FADH 2 , given in equation (4). The anodic current increases with the glucose concentration and the oxidation peak value increases positively from 1.466 μA to 2.641 μA and 1.631 μA to 2.689 μA upon varying concentration from 0-10 mM in the enzymatic and non-enzymatic respectively. The higher electron transfer occurring in the enzymatic reaction leads to the maximum current variation in the enzymatic sensor compared to the non-enzymatic one. GOx reduces flavin adenine dinucleotide (FAD) to FADH 2 which produces electrons and PEG-MnFe 2 O 4 NPs promote the electron transfer as a mediator to the sensing electrode. Higher affinity of PEG-MnFe 2 O 4 NPs towards oxidation of the H 2 O 2 produced during catalytic oxidation of glucose is also responsible for the better performance of enzymatic sensor than the nonenzymatic one as shown in equation (5). Furthermore, the O 2 produced in this reaction can help facilitate the reaction in equation (3). In the non-enzymatic sensor, the mechanism of electro-oxidation of glucose by PEG-MnFe 2 O 4 NPs modified GCE is shown in scheme 1. Glucose in alkaline medium is susceptible to oxidation as glucose forms an enediol structure by losing a proton in the alkaline medium. This enediol structure forms an intermediate complex with the Mn 2+ of PEG-MnFe 2 O 4 leading to the electro-oxidation of glucose which is consistent with the reported literatures [48,49]. The polymer PEG in this reaction acts as a conductive medium which does not affect the electro-oxidation of glucose whereas PEG in the enzymatic sensor aids for the physical adsorption of GOx. All peak current values (I P ) obtained is given in supplementary information tables S1-S3.

Differential pulse voltammetry (DPV)
DPV measurements of varying glucose concentrations in 0.1 M PBS are shown in figure 10. A regular interval pulse applied in system for DPV analysis provides smoother and better graphs. Similar to the above results, oxidation and reduction peaks have exhibited a steady increase in current value as the concentration of glucose is increased. The calibration curve from DPV is also plotted for both the electrodes and the GOx@PEG-MnFe 2 O 4 /GCE has been found to possess a better response to the change in glucose concentration. This response is in accordance with the results obtained in CV, confirming that the enzymatic sensor has better response. Peak current values (I P ) obtained is given in supplementary information table S4. Superior analytical performances in sensitivity and linear range are observed in case of the enzymatic glucose sensor. Glucose oxidase catalyses the oxidation of glucose in the presence of oxygen into D-glucono-1,5-lactone, which then hydrolyzes to gluconic acid and produces H 2 O 2 according to the equations (1)- (3). Since this reaction produces H 2 O 2 which could be further oxidized at the electrode, it leads to an increase in current from the enzymatic sensor. Hence, higher current and sensitivity observed for enzymatic sensor is due to the greater number of electrons involved. Higher affinity of PEG-MnFe 2 O 4 NPs towards oxidation of the H 2 O 2 produced during catalytic oxidation of glucose is also responsible for the better performance of enzymatic sensor [50,51] than the non-enzymatic one. Direct glucose oxidation causes the non-enzymatic sensor to attain the saturation faster leading to the reduced linear range.

Amperometric response
The curve displayed in figure S2 of the enzymatic sensor follows the Michaelis-Menten kinetics. The apparent Michaelis constant (K m app ) and the maximum current (I max ) were obtained and used for analysing the response of the sensor towards glucose. The calibration curve follows a hyperbolic function ( ) / = + y ax b x , where the parameters a and b correspond to the I max and K m app respectively [52,53]. The values of I max and K m app were found to be 3.5 μA and 18.5 mM respectively. The lower K m app value indicates that the enzymatic sensor has greater affinity towards glucose which is comparable to the value reported in literature [54].  The selectivity has been studied as shown in figure 11(c) using the interfering species such as fructose (FR), lactic acid (LA), sucrose (SR), uric acid (UA) and ascorbic acid (AA). As the normal level of glucose in physiological conditions is about 30 times more than these species [55], the selectivity study was conducted in 0.5 mM concentration of these interferants. Enzymatic sensor has exhibited higher selectivity to glucose within the potential range of −0.7057 V to −0.3730 V. Moreover, it exhibited a significant current response towards glucose after glucose addition compared to the negligible response with other interferants. In case of nonenzymatic sensor, selectivity study shows a noticeable current response towards glucose solution containing interference species when compared to enzymatic one. In summary, the enzymatic sensor exhibits high selectivity and specificity to glucose when compared with the non-enzymatic one. The obtained current values are given in supplementary information table S5.

Reproducibility and lifetime analysis
Reproducibility analysis was carried out 5 times with both enzymatic and non-enzymatic glucose sensors for 2 mM glucose concentration in PBS (pH 7.4) keeping the analysis parameters the same. The sensors exhibited reproducibility with a relative standard deviation (R.S.D) of 5.45% and 6.68% for enzymatic and nonenzymatic sensors respectively as shown in figure 12(a). The stability of both sensors was evaluated by monitoring the response current in the presence of 2 mM glucose over 7 days. The relative response of the sensor with respect to the initial value was found to be 95.6% for the enzymatic and 98.7% for the non-enzymatic sensor after 7 days as evident from figure 12(b). Both sensors exhibited substantial stability over the tested period of time owing to the strong binding of enzyme to the conductive polymer layer in the enzymatic sensor and the high stability of the PEG-MnFe 2 O 4 NPs in the non-enzymatic sensor. The comparatively lower response of enzymatic sensor can be attributed to the slight amount of decomposition of GOx from the electrode surface. Hence, these results indicate that both electrodes displayed relatively stable reproducibility and lifetime.

Conclusion
PEG-MnFe 2 O 4 NPs were successfully synthesised and immobilized with GOx for glucose sensing applications. A comparative study of PEG-MnFe 2 O 4 NPs and GOx@PEG-MnFe 2 O 4 has proved better activity for enzymatic sensor due to the presence of GOx which catalyse the glucose oxidation. GOx@PEG-MnFe 2 O 4 (enzymatic sensor) has displayed 1.9 times higher sensitivity with twice the linear range when compared to PEG-MnFe 2 O 4 (non-enzymatic sensor). The better performance exhibited by the enzymatic sensor is due to the electron transfer caused by the catalytic oxidation of glucose by GOx, which is facilitated by the PEG-MnFe 2 O 4 NPs. Higher affinity of these NPs towards oxidation of the H 2 O 2 generated during catalytic oxidation of glucose also contribute towards the enhanced performance. In addition, enzymatic sensor exhibits high selectivity and specificity to glucose within the applied potential range of −0.7057 V to −0.3730 V when compared with the non-enzymatic one. Further, the enzymatic sensor showed significant reproducibility and lifetime due to the stable enzyme immobilization onto the PEG-MnFe 2 O 4 surface. This work emphasises the efficiency of PEG-MnFe 2 O 4 NPs for glucose sensing applications.