ABSTRACT

Since then, different methods have been developed to modify the traditional GOx-CNTs system. Lin et al. described a flow injection amperometric glucose biosensor based on alternatively assembled polyelectrolyte, GOx, and CNTs (PDDA-GOx-PDDA-CNT) [94]. The PDDA-GOx-PDDA sandwich-like structure provided a favorable microenvironment to keep the activity and to prevent the leakage of GOx without the compromise of the excellent electrocatalytic properties of CNTs in the inner layer. Operated at –0.1 V (vs. Ag/AgCl), this glucose biosensor exhibited a wide linear range of 15 µM to 6 mM with a detection limit of 7 µM. Schmidtke et al. developed an innovative method to construct a LBL glucose biosensor using redox polymer [93]. Sodium cholate suspension-dialysis was first applied to produce high-quality dispersion of GOx-SWCNTs composite. Then multilayer films of (PVP-Os)–GOx-SWCNTs were created by repeated, alternating exposure to the redox polymer (PVP-Os) and enzyme (GOx-SWCNTs) solutions. A very high sensitivity of 56 µA·mM-1·cm-2 was achieved ascribed to the combined effect of the LBL composite on electrocatalysis, biocompatibility, and enzyme loading. Compton et al. demonstrated the application of room temperature ionic liquid composite in the CNT-based glucose biosensor [98]. The resultant

GOx-OPFP-MWCNTs composite electrode for the detection of glucose was tested at 0.3 V (vs. SCE) and displayed good sensitivity (2 µM/mM) with the linearity up to 6 mM (without Nafion) or 12 mM (with Nafion). CNTs have also been decorated by metal nanoparticles (NPs) to increase the catalytic performance of both components. Recently, an LBL assembly composed of MWCNTs, Au NPs, and GOx was designed for the specific detection of glucose and exhibited an excellent performance with a wide linear range (0.1-10 mM), good sensitivity (2.53 µA/mM), and a low detection limit (6.7 µM) [84]. Moreover, taking advantage of the synergistic effect of MWCNTs, Au NPs, IL, and CHIT, Lee et al. fabricated an amperometric glucose biosensor integrating these four components on an ITO electrode [130]. The association of Au NPs in the composite electrode greatly improved the sensitivity of the biosensor and negligible interference from uric acid and ascorbic acid was observed. Luong et al. reported the modification of Nafion dissolved SWCNTs with Pt NPs (diameter 2-3 nm) synthesized by wet chemistry [89]. The SWCNT--Pt NPs-GOx composite was deposited on the surface of GCE for the detection of glucose and displayed linearity up to 5 mM, sensitivity of 2.11 µA/mM, and a detection limit of 0.5 µM, which is much lower than that of SWCNTs-GOx-GCE (1 mM) and Pt NPs-GOx-GCE (400 µM). In another research, Xu et al. reported an amperometric glucose biosensor based on Pt NPs electrodeposited on MWCNTs and GOx immobilized in biocompatible CHIT-SiO2 sol-gel [131]. A wide linear range of 1 µM to 23 mM, a short response time within 5 s, and good sensitivity as high as 58.9 µA·mM-1·cm-2 were achieved. Besides Au and Pt, Pd was also applied in the construction of glucose biosensor [132]. Electrodeposition was carried out to co-deposit GOx and Pd NPs onto Nafion-solubilized MWCNTs on GCE. The Pd/MWCNT/GOx greatly enhanced the storage stability and selectivity of the biosensor with linearity up to 12 mM and a detection limit of 0.15 mM. Conducting polymers (CPs) are well known for their capability to increase active surface area and provide good electrical contact between the sensing materials and the electrode [133]. With regard to the application of CPs in CNT-based glucose biosensors, Wang et al. first described the facile preparation route of amperometric enzyme electrode on the basis of the incorporation of MWCNTs and GOx into an electropolymerized polypyrrole (PPy) film [91]. CNTs, herein,

serving as the dopant for maintaining the electrical neutrality during the growth of PPy, retained its electrocatalytic activity and imparted good performance to the glucose biosensor with linearity up to 50 mM and a detection limit of 0.2 mM. Another glucose biosensor was fabricated by loading GOx into polyaniline (PANi)–MWCNTs and Nafion-silica nanocomposite [134]. Cyclic voltammograms of Fe(CN)63-/4-with the as-prepared nanocomposite-modified ITO electrode and a variety of other control electrodes revealed the improved electrocatalytic activity of PANi-MWCNTs. In addition, it has been demonstrated that the highly porous, biocompatible Nafion-silica not only provided a large surface area for the enzyme loading but also maintained 93% of enzyme activity over 20 days. In a recent study, 3,4-ethylenedioxythiophene (EDOT) was electrodeposited on the MWCNT/baked Prussian blue film to effectively entrap GOx [135]. The resultant PEDOT-GOx-baked PB-MWCNTs SPCE showed a linear response to glucose from 1 mM to 10 mM (2.67 µA·mM-1·cm-2) in a flow injection analysis mode at –0.1 V (vs. Ag/AgCl) with highly resolved and reproducible signals (R.S.D.=2.54%). Because of their excellent mechanical strength, CNTs can also be processed to form CNT-based network which can also be applied in the fabrication of biosensors. Xu et al. reported the electrospinning of poly(acrylonitrile-co-acrylic acid) (PANCAA) filled with MWCNTs and covalent binding of GOx through the activation of carboxyl groups on the surface of PANCAA [136]. The nanofibrous membrane was directly deposited on PTE for the amperometric detection of glucose. The water-insoluble polymer (PANCAA) enabled the biosensor to be reused. Furthermore, due to the intrinsic conductivity of CNTs, this type of network can also be used as a flexible, free-standing electrode in the sensing application [137]. A homogeneous SWCNT film was first produced by filtering SWCNT solution through an anodic aluminum oxide membrane and then transfer onto a transparent poly(ethylene terephthalate) support after dissolving Al2O3. The glucose biosensor was realized by encapsulating GOx into the SWCNT film with the aid of Nafion. A linear range of 0.25-3 mM glucose with a detection limit of 97 µM was observed. Last but not least, Fisher et al. developed a Au/Pd nanocube-SWCNT NEE glucose biosensor [138]. GOx was covalently immobilized on Au nanocubes. As mentioned in previous section, the Pd layer provided a seamless connect between SWCNTs and the conductive substrate, as well as a low-resistance contact between SWCNTs and

the Au interface. Meanwhile, Au nanocubes offered the electrode surface ideal enzyme docking ports with excellent biocompatibility. Furthermore, the nanoelectrode environment produced a favorable mass transfer pathway for the access of glucose and the diffusion of H2O2. The resultant biosensor exhibited a wide linear range spanning from 10 μM to 50 mM with a detection limit of 1.3 μM. GOx has also been extensively investigated in DET-based CNT biosensors. As a typical flavoprotein, holo-GOx consists of an apo-GOx and a FAD prosthetic group. The direct electrochemistry of GOx, in fact, is the DET between FAD active center and the electrode surface facilitated by CNTs. The redox reactions of FAD with its natural electron acceptor, oxygen, are shown in the following equations: - ++ + ´ Æ+ Æ 22 2 2 2FAD H FADH (11.2) DETFADH O FAD+H O (11.3) e2 2 Since the electron turnover rate of FAD to electrode in the presence of CNTs is much higher than that of FAD to oxygen, the backward reaction of Eq. 11.2 on the electrode is favored and a well-defined redox peak of FAD can be observed. In a glucose biosensor, with the addition of the analyte, glucose is converted to gluconolactone through following reaction.+ ´ + 2Glucose FAD Gluconolactone FADH (11.4) (11.4) Therefore, the increased concentration of glucose can be correlated with the decreased reduction peak of FAD. Guiseppi-Elie et al. first demonstrated the DET of GOx on SWCNTs [139]. GOx was immobilized on the annealed SWCNT paper and displayed quasi-reversible one-electron transfer process. Herein, SWCNTs were able to “pierce” the glycoprotein shell of GOx and gain access to FAD, which was unapproachable on the smooth surface of traditional electrodes. Afterward, many attempts have been made to achieve DET of GOx by CNTs. For example, the DET of GOx was observed by PSS-MWCNT/Au/IL-modified GCE [140]. A linear range up to 20 mM glucose with a detection limit of 25 μM was obtained. In another research, LBL assembly of GOx/PDDA/SWCNT-modified GCE also displayed a pair of well-behaved redox peaks of FAD and a wide linear range of 1-40 mM for the detection of glucose. Recently, the DET of GOx based on boron-doped CNTs (BCNTs) was reported [141]. As shown in Figure 11.13, the enhanced redox

peaks (already taking the larger background current into account) of BCNTs was due to increased number of defective sites on BCNTs, which was beneficial to enzyme loading and DET between FAD and electrode. This DET-based glucose sensor had a linear range of 0.050.3 mM and a limit of detection of 0.01 mM. The BCNT-modified electrode exhibited good selectivity against the interference from uric acid and ascorbic acid. A similar strategy has been applied to accentuate the DET of FAD using nitrogen-doped CNTs [142]. The resultant glucose biosensor showed a wider linear range (0.02-1.02 mM) with the same detection limit (0.01 mM). These promising results demonstrated the applicability of BCNTs and NCNTs in the fabrication of DET-based, mediator-free, third generation of glucose biosensors [127].