Advanced Biomaterials and Biodevicess

440 Advanced Biomaterials and Biodevices antibody-antigen binding and high sensitivity of electrochemical response, have gained much attention and been applied for clinical immunoassays, not least because this technique combines simple, portable, low-cost elec- trochemical measurement systems. A primary strategy for amplifying the electrochemical signal in immunosensors is associated with a labeled antigen or antibody. Enzyme labeling method is most commonly employed to amplify the number of signal-reporting molecules per biospecific binding between a target biomolecule and an enzyme-labeled biomolecule. An enzyme such as horseradish peroxidase (HRP) or alkaline phosphatase (AP) is usually conjugated to an antibody or an antigen to generate electrochemically detectable species and offered biocatalytic signal amplification in a com- petitive or noncompetitive assay; others such as GOx may be used too (see Figure 12.6) In addition, the label-free configuration has emerged over recent years as an easy and promising method. Below we present some immunosensor devices under these two approaches (enzyme-labeled and label-free) using some common antibodies selective against a variety of antigens, including: α-fetoprotein antigen (anti-AFP), carcinoembryonic antigen (anti-CEA), carbohydrate antigen 19-9 (anti-CA19-9), carcinoma antigen 25 (anti- CA125), and prostate specific antigen (anti-PSA). 12.5.1 α-fetoprotein Antigen α-fetoprotein antigen (AFP) is a glycoprotein with a molecular weight of approximately 70 kDa. It is normally excreted during fetal and neonatal development by the liver, yolk sac, and in small concentrations by the gastrointestinal tract. It is one of the most extensively used clinical tumor markers. The concentration of AFP in healthy human serum is typically below to 25 ng mL-1. Nevertheless, elevated AFP concentration in serum may be an early indication of some cancerous diseases such as hepatocellu- lar cancer, liver metastasis from gastric cancer, testicular cancer and naso- pharyngeal cancer, etc. In 2004, Guan et al. [66] reported one-step immunoassay using PB-modified SPEs. Here anti-human-AFP monoclonal IgG was inmobi- lized onto the electrode and used as the recognition element, and poly- clonal anti-human-AFP IgG labeled with GOx was employed to produce the electrochemical signal. The PB-modified SPE catalyzed H2O2 from the reaction of GOx which allowed quantification of AFP in the sample. Using FIA, the detection range was in the range from 5 to 500 ng mL-1. Finally, the authors compared their results in real serum samples against ELISA. Both

Prussian Blue and Analogues 441 methods were found to give similar results; however, the new approach was less time consuming (30 min) compared to ELISA (2 h). In 2007, Yuan et al. [67] presented a label-free amperometric immuno- sensor for the determination of AFP by immobilizing TiO2 colloids on a PB-modified electrode. AFP responses showed two concentration ranges from 3 to 30 ng mL-1 and from 30 to 300 ng mL-1 with a detection limit of 1 ng mL-1 and exhibited high selectivity, good reproducibility, long-term sta- bility (>2 months) and good repeatability. Finally, the author compared these results, obtained for real serum samples, against chemiluminescence immu- noassays (CLIA) and found good agreement between the two methods. Lately, Hong et al. [68] developed PBNPs and coated them with bovine serum albumin (BSA) to improve their stability. Then gold colloids were loaded on the BSA-coated PBNPs to construct a core-shell-shell nano- structure. Finally, AFP antibody was attached to GNPs and PBNPs/BSA/ GNPs/anti-AFP was used to AFP detection. The dynamic range of the resulting immunosensor for the detection of AFP was from 0.02 to 200 ng mL-1 with a detection limit of 0.006 ng mL-1 and displayed good selectivity, stability and reproducibility. In 2010, Jiang et al. [69] reported an amperometric immunosensor based on the sequential electrodeposition of PB and GNPs on a CNT/GCE surface. Finally, anti-AFP was immobilized onto the GNP surface and BSA employed to block possible remaining active sites of GNP monolayer and avoid any nonspecific adsorption. Under optimal experimental conditions, the immunosensor showed an ultralow limit of detection of 3 pg mL-1 and a linear range from 0.01 to 300 ng mL-1. Moreover, the immunosensor, as well as a commercially available kit, were both used in the determination of AFP in real human serum and showed excellent correlation. Finally, in 2011, Dai et al. [70] developed a sandwich electrochemical immunosensor for the sensitive determination of AFP based on PB-modified hydroxyapatite (PB@HAP) modified with HRP and secondary anti-AFP antibody (Ab2) to fabricate the electrochemical immunosensor label (PB@ HAP/HRP/Ab2). The results indicated that the immunosensor fabricated using PB@HAP/HRP/Ab2 label had high sensitivity, much higher than other labels such as PB@HAP/Ab2, PB/HRP/Ab2 or HAP/HRP/Ab2. Optimized amperometric signals increased linearly with AFP concentration in the range of 0.02 to 8 ng mL-1 with a low detection limit of 9 pg mL-1. 12.5.2 Carcinoembryonic Antigen Carcinoembryonic antigen (CEA) is a glycoprotein involved in cell adhe- sion. It is normally produced during fetal development, but the production

442 Advanced Biomaterials and Biodevices of CEA stops before birth. CEA is one of the most widely used tumor markers worldwide. Its main application is mostly in gastrointestinal can- cers, especially in colorectal malignancy, colorectal carcinoma, pancreatic carcinoma, lung carcinoma and breast carcinoma. In 2009, Ling et al. [71] developed an immunosensor for CEA based on the electrostatic adsorption between the positively charged MnO2 nanoparticles and Chi composite membrane (nano-MnO2/Chi) and the negatively charged PB. Thus, nano-MnO2/Chi membrane was adsorbed onto PB-modified electrode surface and GNPs were electrodeposited to immobilize anti-CEA. Under the optimized conditions, determination of CEA displayed a linear response in two ranges, from 0.25 to 8.0 ng mL-1 and from 8.0 to 100 ng mL-1, with a detection limit of 0.08 ng mL-1. Real serum samples were measured with the immunosensor approach and compared against ELISA, with no significant difference observed between the two methods. Then Zhuo et al. [72] presented a bienzyme functionalized three-layer composite magnetic nanoparticles for electrochemical determination of AFP and CEA. Authors developed a three-layer magnetic nanoparti- cle composed of a Fe3O4 magnetic core, a PB interlayer and a gold shell (Fe3O4/PB/Au). In addition, they used a new signal amplification strategy based on a bienzyme system using HRP and GOx functionalized nanopar- ticles for electrochemical immunosensing using anti-CEA and anti-AFP, respectively. Finally, in 2011, Tang et al. [73] showed the utility of a sensitive elec- trochemical immunoassay for CEA detection with signal dual-amplifica- tion using GOx and PB. The first signal amplification was based on the labeled GOx on the anti-CEA-gold-silver hollow microspheres (anti-CEA/ GOx/Au@Ag@HS) toward the catalytic oxidation of glucose. Thus the enzymatic-generated H2O2 was catalytically reduced by the immobilized PB on the graphene nanosheets. With a sandwich-type immunoassay for- mat, optimized electrochemical immunosensor exhibited a wide dynamic range of 0.005 to 50 ng mL−1 with a low detection limit of 1.0 pg mL−1. Finally, immunosensors were evaluated for clinical serum specimens, and showed good correlation with those obtained by the electrochemilumines- cent method (ECL). 12.5.3 Carbohydrate Antigen 19-9 Carbohydrate antigen 19-9 (CA19-9) is one of the most important car- bohydrate tumor markers associated with biliary, liver, lung and breast cancers, and various other gastro-intestinal malignancies. Furthermore,

Prussian Blue and Analogues 443 CA19-9 is also used as a diagnostic marker for hepatic cyst infection in the autosomal dominant polycystic kidney, and the serum concentrations of CA19-9 increase remarkably in patients with severely damaged renal function In 2010, Liu et al. [74] proposed a novel signal amplification strategy based on PB and Pt hollow nanospheres (Pt/HN) for developing a highly sensitive label-free CA19-9 immunosensor. The authors combined the excellent elec- trocatalytic activity of PB and Pt/HN toward H2O2 reduction to amplify the electrochemical signal. The resulting immunosensors showed a high sensi- tivity and broad linear response to CA19-9 in two ranges from 0.5 to 30 and 30 to 240 U mL-1 with a low detection limit of 0.13 U mL-1. Real samples analyzed with this approach and ELISA methodology revealed good agreement; the authors suggested that the developed immu- noassay may be applied for clinical determination of the CA19-9 level in human serum specimens. 12.5.4 Neuron-specific Enolase Antigen Neuron-specific enolase (NSE) is a glycolytic isoenzyme which is located in central and peripheral neurons and neuroendocrine cells. It has been detected in patients with certain tumors, namely: neuroblastoma, small cell lung cancer, medullary thyroid cancer, carcinoid tumors, pancreatic endocrine tumors, and melanoma. In 2010, Zhong et al. [75] developed a label-free electrochemical immu- nosensor based on PB doped silicon dioxide (SiO2/PB) nanocomposite. SiO2/PB nanocomposite (produced by using a microemulsion method) was used to obtain a nanostructural monolayer on a GCE surface. Next 3-ami- nopropyltriethoxy silane (APTES) was self-assembled in order to obtain an amino-functionalized interface and Chi-stabled gold nanoparticles (Chi/ GNP) were subsequently attached. Finally, anti-NSE was loaded via the adsorption of gold nanoparticles. The immunosensor exhibited good linear behavior in the concentration range from 0.25–5.0 and 5.0–75 ng mL-1 with a limit of detection of 0.1 ng mL-1. The authors demonstrated the application of this proposed immunosensor for the determination of NSE in human serum samples and obtained good results when compared against ELISA 12.5.5 Carcinoma Antigen 125 Cancer antigen 125 or carbohydrate antigen 125 (CA125) is a member of the mucin family glycoproteins expressed in coelomic epithelium during embryonic development. CA-125 has found application as biomarker for

444 Advanced Biomaterials and Biodevices ovarian cancer detection, although it may also be elevated in other can- cers, including endometrial, fallopian tube, lung, breast and gastrointesti- nal cancers. In 2008, Chen et al. [76] immobilized anti-CA125 onto GNPs-modified PBNPs and used this to develop a highly sensitive amperometric immu- nosensor for the detection of CA125. Firstly PBNPs, synthesized using Chi and poly(diallyldimethylammonium chloride) (PDDA) as a protec- tive matrix, were cast onto a GCE surface. Then, GNPs were assembled by the interactions between GNPs and amino groups of Chi and electrostatic interactions between oppositely charged GNPs and PDDA. Finally, anti- CA125 was assembled onto the surface of GNPs. The proposed immu- nosensor showed a high sensitivity, broad linear range and low detection limit for CA125 determination. DPV peak current was proportional to the CA125 concentration in two ranges from 2.0 to 40 and 40 to 100 U mL-1 and the detection limit close to 0.7 U mL-1. Simultaneous analysis of human serum samples with the present approach and with the ELISA pro- tocol suggested acceptable agreement between these two methods. 12.5.6 Human Chorionic Gonadotropin Antigen Human chorionic gonadotropin (HCG) is a hormone produced during pregnancy that is made by the developing placenta after conception, and later by the placental component syncytiotrophoblast. Some cancerous tumors produce this hormone; therefore, elevated levels measured when the patient is not pregnant can signal some type of cancer, such as semi- noma, choriocarcinoma, germ cell tumors, hydatidiform mole formation, teratoma with elements of choriocarcinoma, and islet cell tumor. In 2011, Yang at al. [77] developed an electrochemical immunosensor for HCG based on HRP-functionalized PB-carbon nanotubes-gold nanocom- posites (HRP/GNPs/PB/ CNTs) as labels for signal amplification. Using this approach, Chi hydrogel and TiO2 nanocomposites were first coated onto a GCE surface for the immobilization of primary antibodies (Ab1). Then, the detectable signal was recorded and amplified based on a sandwich-type immu- noassay by the employment of HRP-labeled secondary antibodies (HRP-Ab2) bioconjugate (HRP-Ab2/GNPs/PB/CNTs). Under optimized conditions, the immunosensor showed a linear range from 0.05 to 150 mIU mL-1and a low detection limit of 0.02 mIU mL-1 HCG. The feasibility of the immunoassay for clinical applications was investigated by analyzing several real samples and comparing with the ELISA method. The author found no significant dif- ference between the two methods and anticipated that this immunosensor could be reasonably applied in the clinical determination of HCG.

Prussian Blue and Analogues 445 12.5.7 Prostate Specific Antigen Prostate specific antigen (PSA) is an androgen-regulated serine protease. PSA is secreted by the epithelial cells of the prostate gland. PSA is produced for the ejaculate, where it liquefies semen in the seminal coagulum and allows sperm to swim freely. It is also believed to be instrumental in dissolving cer- vical mucus, allowing the entry of sperm into the uterus. When PSA enters in the circulatory system, it is rapidly bound by protease inhibitors, primarily α1-antichymotrypsin (ACT), although a fraction is inactivated in the lumen by proteolysis and circulates as free PSA (f-PSA). Total PSA (T-PSA) refers to the sum of f-PSA and PSA/ACT complex in serum. T-PSA levels signifi- cantly increase in serum during prostate cancer and other prostatic diseases. Han et al. [78] described in 2012 a sandwich-type electrochemical immunosensor for simultaneous sensitive detection of PSA and fPSA. First, GNPs modified PB and nickel hexacyanoferrates nanoparticles were prepared and used to decorate onion-like mesoporous graphene sheets (O-GS/PBNPs/GNPs and O-GS/NiNPs/GNPs). Then, O-GS/PBNPs/ GNPs and O-GS/NiNPs/GNPs were modified with anti-fPSA and anti- PSA respectively, and streptavidin and biotinylated alkaline phosphatase (bio-AP) were employed to block active sites. Then dual catalysis amplifica- tion was achieved by catalysis of the ascorbic acid 2-phosphate to AA in the presence of bio-AP, and then the enzyme-generated AA was further cata- lytically oxidized by O-GS/PBNPs/GNPs and O-GS/NiNPs/GNPs nano- hybrids at ~0.2 and ~0.4 V vs SCE, respectively. The experiment results showed that the linear range of the proposed immunosensor for simultane- ous determination of fPSA was from 0.02 to 10 ng mL−1 with a detection limit of 7 pg mL−1 and PSA was from 0.01 to 50 ng mL−1 with a detection limit of 3.4 pg mL−1. To evaluate the performance of the proposed immuno- sensor, clinical analysis with serum samples were evaluated and compared against ELISA. The authors found good correlations between those results that confirm the viability of the developed method for real sample assay. 12.5.8 Hepatitis B Antigen Hepatitis B (HB) is an infectious inflammatory illness of the liver caused by the hepatitis B virus (HBV). The virus is transmitted by exposure to infec- tious blood or body fluids such as semen and vaginal fluids, while viral DNA has been detected in the saliva, tears, and urine of chronic carriers. Perinatal infection is a major route of infection; other risk factors for devel- oping HBV infection include working in a healthcare setting, transfusions, dialysis, acupuncture, tattooing, etc.

446 Advanced Biomaterials and Biodevices Recently, He et al. [79] developed a label-free amperometric immuno- sensor based on the electro-deposition of GNPs over PB film. In this way, HB antibody was immobilized onto a modified sensor (GCE/PB/GNPs) surface. Under optimal conditions the peak current response was inversely proportional to the HB antigen concentration. Current changes were pro- portional to HB antigen concentration ranging from 2 to 300 ng mL-1 with a detection limit of 0.4 ng mL-1. Finally, the authors measured and com- pared 50 clinical samples; their results were in good agreement with those using ELISA as a reference method. 12.6 Conclusions In the present chapter we discussed the main advantages of biosensors in biomedical applications. Numerous benefits are driving the replacement by biosensors of other conventional, more sophisticated, analytical methods. Nowadays, biosensors are focused to Point of Care applications, glucose biosensors being the best example due to the serious global problem of diabetes. However, great efforts are being made with a wide range of other illnesses and, in this context, PB and its analogues are promising materi- als due to their electrocatalytic properties. Among different approaches, oxidase-based biosensors are still preferred, and PB is an excellent material for use in the fabrication of biosensors because it acts as an “artificial per- oxidase”. Together with the facile modification of the electrode substrate and the low cost of production, this has led to an on-going replacement of the common enzymatic detection method – horseradish peroxidase (HRP) – which is more expensive and complicated than PB-modified substrates. In addition, over recent years, new applications have appeared in the form of immunosensors, with an important and new creative concept. In this way, novel hybrid configurations have appeared, using PB and analogues as well as other modern materials such as graphene, carbon nanotubes, mag- netic beads, gold nanoparticles, etc. Based on these results, it is clear that during the next decades, PB and its analogues will have important roles in the future development of biomedical devices for health care. Acknowledgment The funds for the development of this contribution have been provided by subprograma INNCORPORA TU (INC–TU–2011–1621) from Ministerio de Economía y Competitividad, Ministerio de Industria, Turismo y Comercio

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13 Efficiency of Biosensors as New Generation of Analytical Approaches at the Biochemical Diagnostics of Diseases N.F. Starodub* and M. D. Melnychuk National University of Life and Environmental Sciences of Ukraine, Kiev, Ukraine Abstract The laboratory diagnostics is based on the qualitative and quantitative determina- tion of a number of biochemical parameters in biological fluids. Among recent widespread methods based on modern instrumental electronic devices the biosen- sors are known very well already. Today a lot of types of biosensors are developed by the realization of different physical and chemical principles for the registra- tion of the specific generated signals. There a great significance plays the choosing effective ways of the modification of biosensor surface, optimization of the condi- tions for the immobilization of biological material and choosing the most available protocol of analysis. In this chapter it was demonstrated the practical realization of the biosen- soric principles (using the optical and electrochemical transducers) at the simple, cheapness, express, selective and sensitive analysis of such biochemical quantities as: glucose, cholesterol, urea, sodium, potassium, calcium, certain immune com- ponents and etc. At the achievement of better parameters of biosensor work there is possibility to obtain the sensitivity analysis which exceeds that which may be provided by the traditionally used methods. Moreover biosensors may give advan- tages before traditional methods in respect of time which should be spend for analysis. Device based on the biosensor principle can be as portable. The possibil- ity to apply artificial substances, in particular, using calyx[4]arenas, some template *Corresponding author: [email protected] Ashutosh Tiwari and Anis N. Nordin (eds.) Advanced Biomaterials and Biodevices, (451–486) 2014 © Scrivener Publishing LLC 451

452 Advanced Biomaterials and Biodevices structures as selective sites and the development of multi-parametrical and multi- functional biosensors are considered as further perspective of the creation of the instrumental analytical devices of new generation. The very important role belongs to the use of nano-structured materials, nano-particles and nano-tubes in form of new perspective transducers. Keywords: Biosensors, transducers, selective sites, types, diagnostics, biochemi- cal quantities, determination 13.1 Introduction Laboratory diagnostics as part of the clinical biochemistry is based on the qualitative and quantitative determination of a number of biochemi- cal parameters in biological fluids. The field of its activity includes exam- ining the nature of the changes of biochemical quantities at a number of physiological and pathological conditions and the development of meth- ods for their determination. Among recent widespread methods based on the modern instrumental electronic devices the biosensors are known very well already. Today, these devices are used in the clinical diagnos- tics to determine such important parameters as glucose, cholesterol, urea, sodium, potassium, calcium, certain immune components and etc. 13.2 General Approaches for the Development of Optical Immune Biosensors This portion, considers aspects of the application of the biosensor based on the fiber optics and realized principles of chemiluminescence (ChL), fluorescence, non-emitting energy transfer and “evenescens” wave as well as uses of surface plasmon resonance (SPR) and nano-structured porous silicin. 13.2.1 Fiber Optic Immune Biosensors for Diagnostics Currently, the immunochemical analysis with the label in form of enzyme, chemo-, bioluminescent or fluorescent substances is widely used. No doubt, further progress in this field of medicine will be made by the devel- opment of the immune biosensors providing analysis of an even higher sensitivity, expressive, simplicity, used inexpensive equipment and requir- ing minimal sample pretreatment.

Biosensors for Biochemical Diagnostics of Diseases 453 13.2.1.1 Construction, Fluorescent Labels, Used Reagents and Measurements [1] An argon laser was served as the source of light. The optrodes (diame- ter: 1, length: 50 mm) were made from mono fiber light conductors. They were pretreated with a mixture containing concentrated hydrochloric acid and dichloroethane in equal doses. Before immobilization the surface of optrodes was activated with cyanogen bromide at -150 0C in the presence of triethylamine. Then they were immersed into the antigen (Ag) solution for about 30 min and in 0.1 M glycine solution for 40 min to block the remaining free functional groups. For the immobilization of Ab to human IgG the optrodes were aminosilanised and activated with glutaraldehyde (GA). After such treatment they were immersed into the conjugate solu- tion (1:10 titer) prepared in phosphate buffer (PhB), pH 7.2 for 1 h at room temperature. The remaining free functional groups were blocked as men- tioned before. The prepared optrodes were stored dry and sterile at 4 0C. To measure the level of non-specific IgG using the fiber optic fluorescent biosensor based on the principle non-emitting energy transfer, one half of the prepared optrodes was placed in the solution containing human IgG (1:20 conjugate titer). This IgG were labeled by tetramethylrhodamine iso- thiocyanates (TRIC). The other ones were introduced in the same solu- tion but with the B-phycoerythryn (B-Phyco)-labeled IgG. Both labeled conjugates were dissolved in PBS, pH 7.2 containing tween-20 and bovine serum albumin (BSA) in final concentration of 0.05 and l%, respectively (buffer A). The degree of antibody (Ab) saturation with Ag was estimated by measuring the value of the signal B = If,/Ifd, where If, is maximum emis- sion intensity at band-pass acceptor fluorescence (TRIC, 560 and B-Phy, 575 nm); Ifd is a maximum emission intensity at band-pass donor excita- tion (TRIC, 488 nm). Every other 3 min during the incubation one of the optrodes was taken out to measure the fluorescence signal. The minimum saturation period was 21 min. For sensor calibration the optrodes were immersed one by one into solutions containing a number of consecutive dilutions of human IgG in buffer A starting from a concentration of 500 ng/ml. While being measured the optrode was placed into a black Teflon microcell. The block-scheme of such an immunosensor is given in Figure 13.1 [A]. The optrode was connected to the beam splitter by a special joint. The laser emission spread through one beam splitter branch and the fluorescence emission to the photo detector through the other one. When the complex of immobilized labeled Ab (donor) -1abelled Ag (acceptor) is formed, the fluorescent labels approach each other thus provoking an energy transfer,

454 Advanced Biomaterials and Biodevices Figure 13.1 [A] block-scheme of the sensor: 1– laser: 2– beamsplitter, 3 – optrode, 4 – clamp; 5 – lens, 6 – interference filter, 7 – shutter, 8 – photoelectric cell, 9 – photon counter, 10 – flat parallel plate, 11– light fiber, 12 – light diode, 13 – amplifier, 14 – voltmeter, 15 – recorder, 16 – casing, 17– capillary. [B]: principle of the analysis fulfillment with the help of the optical immune biosensor based on the non-emitting energy transfer, where: 1 – optrod, 2 – sample, 3 – Ab immobilized and labeled by donor fluorochrome, 4 – Ag to be analyzed, 5 – standard Ag labeled by acceptor fluorochrome. the emission of which is registered. To remove artifact’s caused by instabil- ity of the light source and other factors affecting donor fluorescence the signal value was expressed as the ratio of the level of band-pass acceptor fluorescence and band-pass donor excitation. It was performed by the introducing another semi-transparent mirror (10) (at 450 angles to the laser ray), optic fiber (11) and photodiode (12). A special device for replac- ing interference filters was installed in front of the photoelectric cell to ensure it operated with two fluorescence labels. Theoretically the idea of an immune biosensor based on energy transfer may be expressed as follows: [Ab’] + [Ab*] = [Ab’-Ag*] and [Ab’-Ag*] + [Ag] = [Ab’-Ag] + [Ag*], where Ab’ – immobilized Ab labeled with donor fluoroforme, Ag* – antigen labeled with acceptor fluoroforme, Ab’ -Ag* – the complex of immobilized labeled antibody with labeled antigen, Ab’ -Ag – the complex of immobilized labeled antibody with analyzed antigen (Figure 13.1 [B]). Since the amount of the immobilized Ab is constant the fluorescence intensity is inversely proportional to the Ag concentration. There is no need to separate the immune complex from individual components. For the calibration it was used 20 identical optrodes after their immersing into the solutions containing consecutive four-fold serum dilutions (starting from 1:20 M), standard titer of the monoclonal Ab labeled by horse radish peroxidase (HRP). Optrodes with immobilized BSA were used as controls. In parallel the ELISA- method was performed [2]. The activity of HRP was

Biosensors for Biochemical Diagnostics of Diseases 455 tested by enhanced ChL in special mixture containing H2O2, p-iodinephe- nol, luminol,tris-HCl buffer (TB), pH 8.1. 13.2.1.2 Control of the Total IgG Content by the Direct Determination of Fluorescent Label [1] It was stated that in case of the determination of total human IgG the sen- sitivity of biosensor using fluorescein isothiocyanate (FITC) and B-Phyco is an order of magnitude higher than that based on FITC-TRIC. This dif- ference is result from the following factors: (i) the quantum output for B-Phyco is much higher than that for TRIC; (ii) the absorption spectrum of B-Phyco has a greater overlap with FITC as compared to that of TRIC; (iii) the stock width of B-Phyco is much larger than that of TRIC. When using biosensors of this type the measurement may be conducted in a real time regime. Adding 0.01% polyethyleneglycol 6000 (PEG) to the solution reduces the response time by half possible due to fastened Ag-Ab interac- tion triggered by destruction of the ion screen surrounding the proteins. When the system of fluorescence labels FITC-TRIC is used there is pos- sible to register about 10 ng/ml of IgG. About 80% of the maximum pos- sible signal may be realized through 2 min after beginning incubation. The obtained data correspond to the most sensitive ELISA-method. However, when comparing its timeframe (3 h and more) and biosensor analysis time (from 15 min to 2 h), the latter may be considered more preferable for bio- chemical diagnostics. In analysis of the IgG of the same concentration a 4% variation for five optrodes was observed. To study the possibilities of reuse, optrodes were immersed into 0.1 M glycine-HCl buffer (GB), pH 2.5 for 10 min and washed with buffer A. The sensitivity analysis was somewhat lowered. After 5 consecutive regenerations it was reduced two-fold. This may suggest that the primary conformation of the immobilized Ab can be destroyed by substances dissociated Ag-Ab complex. Therefore, replace- able optrodes prepared beforehand are recommended. To improve proce- dure renew of optrodes for their many single using we have tried to apply different types of buffer systems, urea, dithiotriethol, sodium dodecylsul- fate and others. The best finding was obtained in the case with the solution contained 6 M urea, pH 7.4. It allows about 20–30 consecutive measure- ments to be made with a final signal failure of 20%. Thus, the developed opto-immune biosensor is of lucid technology and design, combining effective methods of immobilization of the immuno- logic reaction component and original methods of immune complex for- mation. Such biosensors are highly sensitive and produce a fast response. They may be recommended for the wide use in on line and in situ analysis

456 Advanced Biomaterials and Biodevices when studying the biochemical parameters of an organism which is tre- mendously important in medical practice. 13.2.2 Fiber Pptic Immune Biosensor Based on the Principle of the “Evanescent” Wave The principle of non-emissive energy transfer between two fluorescent labels, when one of them is belonging to Ag and the other one to Ab, may be used for creating optical immune biosensors [3, 4]. 13.2.2.1 Construction, Used Reagents and Measurements [5] A schema for the proposed biosensor is given in Figure 13.2. Light from the argon laser (1) spreads through the flat parallel plate (2) and semitrans- parent mirror (9), and then falls on the entrance butt-end of the optrode (10). The fluorescent signal from the lateral optrode surface caused by the evanescent wave was tunneled through the back of the optrode. This sig- nal is reflected from the semi-transparent mirror (9), falling on the higher effective beam splitter. One part of the rays (reflected light) is directed to the photoelectric cell with a maximum band-pass at 575 nm (semi wide 8 nm). The other part of the rays (passed through 11) falls on the second Figure 13.2 A. Scheme of optical immune biosensor based on the principle of the “evanescent” waive for the simultaneous measurements of the presence of two analytes in the solution (detailed explanation is in the text). B. Overall view of competitve principle of analysis. Where: 1 – optrode, 2 – sample; 3, 4 – immobilized Ab to phenatoine and lidocaine, respectively, 5 – phenatoine, 6 – phenatoine labeled by FITC, 7 – lidocaine, 8 – lidocaine labeled by R-Phy, 9-11 – fluorescent irradiation that arouses at the excitation FITC and R-Phy, respectively (B).

Biosensors for Biochemical Diagnostics of Diseases 457 photoelectric cell through an interference filter with maximum band-pass at 520 nm (12). The signal from each photoelectric cell (15) enters the two- channel photon counter and is then passed to the two-channel recorder (16) through the analogous transformer. To remove artifacts caused by instability of the light source, part of the laser rays reflected on the flat parallel plate is arranged by lenses (3) and with the help of fiber light con- ductors (4) reaches the photodiode (5). The photodiode signal is amplified and registered by a voltmeter (7) and recorder (8). The value of the specific signal (B) is calculated as follows: B = (I –I0)/I, where I – maximum value of the signal, I0 – the light background of pho- toelectric cell. An optrode is fixed at the distal end only. It permits rapid replacement of them and maximum use of the energy of the evanescent wave. The 1.6-mm inner diameter of the microcell provides a usable capac- ity of up to 20 μl. For such a biosensor, we used an argon laser as the light source as well as the optrodes from quartz optic fibers (diameter 1.3 mm, length 60 mm). In experiments we have taken phenatoine, lidocaine, spe- cific Ab to them and as well as they labeled forms by FITC. All operations were carried out at the room temperature. The optic fibers were pretreated as follows: • First, to form a high concentration of hydroxylic groups on the surface the optrodes were immersed in hot acetone for 5 min. Then, these glass fibers were boiled in water for 5 h. To obtain uniformly distributed active groups on the surface the optrodes were treated with 2% monofunctional 4-aminobutyldimethylmethoxysilane in acetone for 40 min at 60 0C. After washing with distilled water the optrode sur- face was activated by 5% GA at pH 6.8 for 40 min. Then the optrodes were washed carefully in 0.05 M PhB, pH 7.2. For immobilization of the Ab the optrodes with the modified surface were immersed for about 1 h into a solution contain- ing 500 ng/ml of the mixture (1:l) of the Ab to phenatoine and lidocaine prepared in the PhB mentioned above. The remaining free functional dialdehyde groups were blocked by a 0.1 M glycine solution for 40 min. Finally, the optrodes were washed with 0.05 M PhB, pH 7.2 and dried in air. The concentrations of the determined substances are estimated by the competitive method. In this case a sample containing equal volumes of control and analyzed solution is introduced into the microcell. As refer- ence samples we used phenatoine labeled with FITC and lidocaine labeled with B-Phyco which were diluted 1:20. The solution for the calibration

458 Advanced Biomaterials and Biodevices included phenatoine and lidocaine at the various dilutions starting from a concentration of 200 μg/ml and 50 mM PhB, pH 7.2, containing 1% BSA and 0.05% tween-20 was used in all cases. 13.2.2.2 Determination of Some Pharmaceutical Substances [5] The minimal biosensor’s response time in the concentrations range of 20 ng-20 mg/ml was 50 and 75 sec for lidocaine and phenatoine, respec- tively. Although the measured concentration is equal and the kinetic curves similar to Langmuir’s adsorption curves, there was a difference in kinetic response of the above mentioned substances. Firstly, it is possible since the detected substances with like diffusion factors have different association constants in the immunochemical reaction. Secondly, FITC and B-Phyco have different stocks width (30 and 85 nm, respectively) which is an impor- tant index for calculating the signal/noise ratio. The biosensor’s maximum response was achieved through 2 min at 100 ng/ml. This is much less than that required for a therapeutic effect. To control the selectivity of the analysis, such substances as digoxin, fer- ritin, human Ig, phenobarbital, methotrexate, digitoxin, gentamicin and theophylline were used. The response value of the biosensor to the above- mentioned list of substances was no more than 6.7% of the specific sig- nal. When analyzing the same concentration of phenatoine or lidocaine, a 3–8% variation was observed among five optrodes. After 2 months of storing optrodes in dry and sterile conditions at 4 0C the response time was increased by 15 s, while sensitivity decreased 2%. Such a biosensor may find wide application in different fields, especially medicine. It is highly selective and sensitive with the response time in sev- eral minutes that allows fulfilling analysis of many pharmaceutical sub- stances online or under in situ regimes. 13.2.3 Immune Biosensor Based on the Effect of the Enhanced Chemiluminescence (ChL) [6] This approach may also be effective for medical diagnostics as recently some variants of such analysis in form of the ‘dot’-ELISA-method were developed [7, 8]. 13.2.3.1 Construction of Biosensor, Used Reagents and Measurements A scheme for such biosensor is presented in Figure 13.3. The optrode with the immobilized material was placed into a teflon cell of 8 μl capacity

Biosensors for Biochemical Diagnostics of Diseases 459 Figrue 13.3 Scheme of the sensor (A) and sensitive element (B): I – capillary, 2 – optrode, 3 – shutter, 4 – light diode, 5 – photon counter, 6 – recorder, 7 – light fiber, 8 – biological material or special membrane, 9 – capacitor, 10 – casing. containing specific Ab of known dilution and Ab labeled by the HRP. The inoculated and washed optrode was placed in the cell with containing the substrate for enhanced ChL. The signal resulting from the agitated luminol passed to the photon meter and recorder. Optrodes were pretreated with a mixture which contained con- centrated hydrochloric acid and dichloroethane in equal doses. One approach was achieved by the direct immobilization of the biological material on the surface of the optrode and on the inside wall of the spe- cial capillary. In the latter case, a special membrane with a tightly bound optical fibre was used. Direct immobilization of the biological material was fulfilled in two ways. In one case before, the immobilization the optrode surface was acti- vated with cyanogen bromide at -150 0C in the presence of triethylamine. Then the optrodes were immersed in Ag or Ab solution for about 30 min and in 0.1 M glycine solution for 40 min to block the remaining free func- tional groups. In the other case, the optrodes were previously silanized and then activated with GA. The process of silanization was accomplished as follows: • aminopropyltriethoxysilane (APTES) was frozen and situ- ated in a vacuum, where the optrodes were placed; • the vacuum container was filled with APTES vapor in which the optrodes were held for no less than 12 h; • the container was then placed for 3 h in vapors of GA obtained by the method mentioned previously. Interaction of the optrode surface with the biological materials, and

460 Advanced Biomaterials and Biodevices blocking of the remaining free functional groups, were car- ried out as per the first method. If BrCN was used, the bio- logical material was bound 1.8 times more effectively. In this case, the signal was higher and consequently the method became more precise. Nitrocellulose (NC) was used for the membrane strips, which were immersed in the biomaterial solution (concentration of 0.1 μg/ml, 1 h at room tempera- ture). Alter that they were desiccated and the remaining free functional groups were blocked. This procedure was similar to that used in the “dot”-ELISA-method. Careful examination disclosed that the biomaterial may be either directly immobilized on the surface of the optrode, or tightly connected with the special membrane. One of the components of the immunochemi- cal reaction was conjugated with HRP. The latter catalysis breaks up H2O2 into oxygen radicals which are oxidized by luminol to yield aminophthalic acid. The excited part of the oxidized molecules of luminol emits quanta of light. The number of flashes may be drastically increased in the pres- ence of some chemical compounds. An example of such an enhancer is used p-iodophenol. The level of ChL depends upon a number of factors: concentration of H2O2, p-iodophenol, luminol, pH media etc. The optimal conditions were: 2 mM H2O2, 0.06 mM p-iodophenol, 0.06 mM luminol and 50 mM TB, pH 8.1. Immunochemical reactions were carried out in two different regimes: (i) in stationary conditions, when the optrode was placed in the measur- ing cell or in a capillary into which the analyzing solution was sucked and on the wall of which the immunochemical interaction would occur; (ii) in flowing regimes, where the analyzing solution was pumped through the cell. The optrode alone or placed in a capillary was immersed in a special non-transparent cuvette. The value of the luminescent signal was mea- sured for a few minutes. 13.2.3.2 Determination of the Individual Indexes which are Important for Diagnostics On the surface of the optrode or the NC membrane, IgG was immobilized. The measuring cell was filled with a solution of Ab to IgG, which were con- jugated with HRP, and a sample of diluted serum. The titer of Ab labeled by HRP was 1:500. All dilutions were made in PBS, pH 7.2, containing 0.05% twin-20 and 1% BSA. After 10 minutes, the cell was washed and refilled with a solution of the components needed to carry out the reaction

Biosensors for Biochemical Diagnostics of Diseases 461 of enhanced ChL. Two-fold dilutions of IgG starting from 300 ng/ml were used as standard solutions. The signal reaches maximum value after 17 min of exposure in the solution to be analyzed. The response period was half as much after the addition of PEG in 0.01% final concentration. The sensitivity of analysis by the biosensors was similar with the ELISA method and achieved about several ng/ml. In case of the determination of the chorionic honadotropine (ChH) level the specific monoclonal Ab to its α-subunit were immobilized on the surface of the optrodes. Then they were immersed into cell which contained 5 μl of blood serum diluted in PBS (1:3) in the presence of 0.05% twin-20 and 1% BSA as well as 5 μl of specific Ab conjugated with HRP (dilution 1:100). For the calibration curve the sample contained 5 μl of the above mentioned mixture and 5 μl of consecutive two-fold ChH dilutions (starting from 150 ng/ml) were used. The sensitivity of the analysis reached 10 ng/ml and the response time was 5–7 min. These results were confirmed by the ELISA-method in the range of concentration of 10–100 ng/ml. At the control of the estradiol-17 concentration the specific monoclonal Ab were covalently bound to the optrode surface. Then it was immersed (on 10 min) in a cell containing 0.01 ml of human blood serum diluted with PBS (1:3) in the presence of 0.05% of twin-20. Measurements were carried out by two different ways. In the first one after washing in the above mentioned buffer the optrodes were placed in a cell containing 0.01 ml of estradiol-isoluminol conjugate (50 ng/ml). Then they were placed in a non-transparent cell filled with a mixture of p-iodophenol (0.06 mM), HRP (4  mM), H2O2 (20 mM) and TB (10 mM), pH 8.2. For calibartion of optrodes the samples containing 5 μl of consecutive two-fold estra- diol-17 dilutions starting from 50 ng/ml and 5 μl (50 ng/ml) of estradiol- isoluminol conjugate were used. According to the second way (sandwich variant) the estradiol-isoluminol conjugate was replaced by specific Ab linked with HRP. The controlled concentration was the same as in case for ChH. In both cases the sensitivity was about 6 ng/ml. The total time of analysis in the second case was approximately in two times less than in the first one (~20 min). A strong correlation (r=0.95) was found between these results and obtained by the ELISA- method. In the next experiments, the developed method was examined at the con- trol of α-2-interferon in the serum blood. The monoclonal Ab to recombi- nant α-2-mterferon were immobilized on the inside wall of a capillary (inner diameter 1.0, length 6 mm). The quartz optical fiber (diameter 0.8, length 75 mm) was put into the capillary. The sample (2 μl) of purified α-2-interferon linked with HRP (1:200 dilutions with 0.05% twin-20 and 1% BSA) and the analyzing sample (of the same volume) were sucked into the capillary. After

462 Advanced Biomaterials and Biodevices 10–15 min the capillary was washed and refilled with the solution for deter- mination of activity by enhanced ChL. For obtaining the calibration curves the samples were consecutive two-fold dilututed starting at 300 ng/ml. The sensitivity of analysis was roughly 9.5 ng/ml and the total time was less than 20 min. The range of reliable results was from 20 to 150 ng/ml. This approach was used too for the control of Salmonella typhimurium level in solution. About 1 ml of its suspension was filtered through an NC filter. After desiccation it was washed (PBS containing 10% of BSA). Then it was soaked (15 min) in the mixture of monoclonal Ab to receptor linked with HRP and prepared on PBS, pH 7.2 containing 5% BSA and 0.5% twin- 20. Finally the washed filter was fixed at the top of the optrode and placed in a non-transparent cell containing the necessary components for ChL. The sensitivity of analysis was on the level of 50 cells/1 ml that was slightly better than with the ELISA-method. In case of revealing Ab specific to influenza virus (Singapore 222/79) the last in inactive form was immobilized on the inside wall of the reactor (capillary). First 0.1 ml of the mixture of serum blood diluted in PBS in the presence of 0.05% of twin-20 with BSA and then 0.1 ml of the protein A linked with HRP, dilution 1:100, were passed through the reactor. After reactor was washed and the optical fiber has been immersed, 0.2 ml of the solution for the reaction of enhanced ChL was passed. The titer of Ab determined was 1:9152. This is slightly less than that obtained by the ELISA-method. However, the total time of analysis was only 30 min in com- parison with minimum 3 h needed for the ELISA-method. The signal level of the biosensor depended upon the current velocity of the measured sample and the components of the solution for enhanced ChL. Undoubtedly, if these parameters are optimized, the measurements will be more precise. The best results at the reusing optrodes were obtained in the case of using 6 M urea, pH 7.4 as it was shown above. Regeneration allows about 20–30 consecutive measurements with a final signal failure at 20%. The main advantages of these biosensors in comparison with the ELISA- method are low overall analysis time and the simplicity. Such sensors are very effective for making separate analyses. We believe that in the near future thte will be widely used not only in medicine but also in others field. 13.2.4 Immune Biosensor Based on the Photoluminescence (PhL) of Porous Silicon (PS) [9–17] With the discovery of enhanced photo- and electroluminescence of PS [18] numerous investigations have been undertaken to study these effects and use them at the creation of optoelectronic devices such as light emit- ting devices, gas sensors, photodetectors and solar cells [19, 20]. Being a

Biosensors for Biochemical Diagnostics of Diseases 463 promising material for the technology, PS also excites great interest among scientists working with biosensors aimed to detect a biological substance of quickly and in small quantities. This is critically important for the express diagnostics of diseases and environmental monitoring. Several methods have been proposed to obtain PS and to use its properties at the creation of different devices [9–20]. Earlier [10–16] we have informed about the development of a biosensor based on the PS PhL. It was demonstrated too [9] the possibility using for the creation of an biosensor for quick detection of myoglobin (Mb) as one of markers of the heart failure [9, 21]. 13.2.4.1 Sensor Construction, Preparation of PS Samples and Reactive Used [9] The PS samples with the size about 4×4 mm2, were obtained from mono- crystalline silicon of p-type (the resistance about 10 ῼ/cm). They were treated by laser beam YAG:Nd-laser (with the λ 1.06 mm; impulse dura- tion t=5×10–4c; E=0.5J) and then chemically etched in a solution of HF:HNO:H2O (volume ration 1:3:5) during 5–40 min at the room tem- perature. Then silicon was passivated. PS was characterized by the size of pores equal to 10–100 nm. PS PhL was excited by He–Cd laser (with λ 440 nm, P = 0.001 W) and measured by the standard monochromator. The visible PhL had λ equal to 650 nm and its time decay was described by “stretch” exponent. Samples of PS were washed several times with distilled water, then with absolute alcohol and at last they were dried at room tem- perature (Figure 13.4). (A) Monocrystaline Silicon (B) 1 (p-type, ρ = 10 x cm) 9 2 Laser Irradiation 3 (YAG: Nd laser; λ=1.06 m; τ = 2x10–4c; E = 0.3 J) 4 Chemical Treatment 5 (HF : HNO4: H2O = 1 : 3 : 5), 3–10 6 Porous Silicon 7 (sample size: 4x4 mm2; pore diameter: 50–400 nm) Photoluminescence Recording (He-Cd laser; λ=440 nm; P = 0.001 W) Figure 13.4 Block-schemes of PS fabrication (A) and the set-up for PS PL measurement (B) where: 1 – He-Cd laser and N2-1aser, 2 – modulator, 3 – prism, 4 – mirror, 5 – quartz cell, 6 – PS-sample, 7 – spectrophotometer, 8 – photomultiplier, 9 – computer.

464 Advanced Biomaterials and Biodevices The mouse monoclonal Ab to anti-human Mb were deposited on the surface of PS samples by spontaneous sorption. In this case they were immersed into the Ab solution prepared on 20 mM PBS, pH 7.3. As ade- quate time for further experiments it was chosen period in 60 min since the time in 15 and 30 min of Ab immobilization are rather short to have on the surface such a quantity of Ab which is sufficient for the sensitive detec- tion of Mb. Then PS surface was washed with PBS, pH 7.3. Three samples of human serum were used: undiluted, diluted to 1:10 and to 1:100. Mb concentration was determinhed by the ELISA-method too. In last case monoclonal mouse Ab to two different Ag sites of Mb were used. One type of the Ab was immobilized on the surface of plates, while other ones, con- jugated with HRP, served to detect Mb. 13.2.4.2 Study of the Main Characteristics of PS PhL and Quantitative Mb Determination [9, 10] Earlier, it was demonstrated [10] that at the contact of PS with distilled water, or 10 mM TB, pH 7.3, or TBS, or solutions of BSA, or rabbit IgG, or sheep anti-rabbit IgG prepared in above mentioned buffers, its typi- cal PhL spectra remained almost immutable during no less than 2–2.5 h. At the same time after previous deposition of Ab or Ag on the PS and their consecutive contact with the solution of corresponding Ag or Ab a great reduction of the PhL intensity was observed. According to the exist- ing ideas [11–17] the nature of the visible PhL of PS can be explained by a process of dehydrogenization of the PS surface, which takes place after a specific immune complex formation. Hydrogen is released from silicon bonds and subsequently captures by the immune complex. Torn Si bonds are known to intensify non-radioactive channel of recombination, which leads to the decrease of the PhL intensity. The overall view of PS by atom force microscopy in different situation is shown in Figure 13.5. In the following experiments an adequate concentration of Ab to be deposited on the PS surface was studied. PS samples were immersed into the Ab solutions of different concentrations for 60 min. Then PS samples with Ab on the surface were immersed into 20 mM Na-PBS (pH 7.3) with Mb in the concentration equal to 0.1 μg/ml. It was shown that after the first 10 min of the experiments in all samples the decrease of the PhL inten- sity already took place. In the case of the PS samples on which Ab were deposited in the solutions of 20 μg/ml, through 30–35 min after the begin- ning of the measurements the decrease in the PhL intensity was up to 80% from the initial level. In 20–30 min the magnitude of the PhL intensity was 50% from the initial level. During 30–35 min the PhL intensity of PS

Biosensors for Biochemical Diagnostics of Diseases 465 12 3 MKM 2,0 0,5 0,4 0,5 0,4 0,4 1,5 1,0 0,3 1,0 0,3 0,3 1,0 1,5 0,2 1,5 0,2 0,5 0,1 0,1 0,2 0,1 Figure 13.5 Porous silicon at the atom force microscopy. 1 – bare surface, 2 – with specific antibodies, 3 – with Ab-Ag complex. samples with Ab immobilized from the solution of 4 μg/ml decreased only for 40%, while in the case of Ab concentrations 100 and 500 μg/ml the changes in the PhL intensity achieved their maximum level. At the same time, the concentrations equal to 100 and 500 μg/ml were rather similar, which allowed considering that the concentration of Ab equal to 100 μg/ ml was sufficient for efficient registration of Mb. It was shown that the PhL intensity did not change at the immersion of the PS sample with deposited Ab (concentration of Ab: 100 μg/ml; time of immobilisation: 60 min) into the Mb solution of the concentration equal to 1.0 ng/ml during 15 min. The magnitude of concentrations detected lied in the range from 0.01 to 10 μg of Mb in 1ml of solution. PhL intensity of PS samples retained in undiluted serum was about 50%, while PhL intensity of PS samples immersed into diluted serum samples approached to 100%. At the same time a difference between data obtained for undiluted serum and Mb solution 100 ng/ml was observed. In case of Mb solution the PhL intensity decreased for 40% while for undiluted serum this value was 50%. Such a difference in the magnitude of the decrease in the PhL intensity can testify about probable influence of other PS samples with Ab on the surfaces (concentration of Ab: 100 μg/ml; time of immobi- lization: 60 min) were immersed into those three samples of serum for 15 min and then the PhL intensity was measured. The results showed that the PhL intensity of PS samples retained in undiluted serum was about 50%, while PhL intensity of PS samples immersed into diluted serum samples approached to 100%. Such changes of the PhL intensity in case of undi- luted serum can be explained either by the fact that serum of healthy peo- ple can contain Mb in concentrations up to 100 ng/ml or by the influence of others proteins. The last is less probable, since according to our previous results it was shown that sorption of albumin or IgG on the PS surface did not affect the PhL intensity. At the same time, a difference between data

466 Advanced Biomaterials and Biodevices obtained for undiluted serum and Mb solution 100 ng/ml. was observed. In order to check such an influence the following experiments were car- ried out. Different quantities of Mb were added to the samples of serum diluted by 20 mM Na-PBS (pH 7.3) till 1:10 and then solutions obtained were used to measure the changes of the PhL intensity during 15 min. The results were a little bit different from that which was obtained for the Mb solution. This difference can be explained by the presence of initial quanti- ties of Mb in the serum. To get an idea about the correlation of the data obtained with the results showed by the ELISA-method the following mea- surements were undertaken. Samples of serum (dilution 1:10) from three patients, suffering from heart failure disease of different degree of serious- ness, were taken for analysis by both methods. Data for Mb concentra- tion, determined with the help of the sensor, were obtained according to a calibration curve constructed in advance. Since usual concentration of Mb in the serum of healthy people is about 0.1 μg/ml, and at the heart failure disease it can increase up to 1 μg/ml, the range of concentrations between 0.1 μg/ml and 1 μg/ml was used for the construction of calibration curve. The data obtained with the help of both methods had difference no more than 5%. Since the durability of the analysis by the ELISA-method is no less than 3 h. and the overall time of measurements by the biosensor about 15 min, it is possible to assume that the last has high potential for the fur- ther application in the area of medical diagnostics. 13.2.5 Direct Electrometric Approach to Register Interaction Between Biological Molecules [18, 19] To completely fulfill all practice demands in respect of simplicity and cheapness of the analysis it was studies capability of the electrometrical method for the providing very sensitive registration of the changes of physical-chemical properties of the test structures [22–25], seems to be more expedient, especially if it is concerned for the materials with the high capability of sorption to the surface. The metal-semiconductor specimens were prepared by vacuum beam- evaporation of Ni on a surface of silicon with using the standard wafers of technology. The thickness of the films was equal to: 50–300 A0. The surface- barrier structures (SBS) with a super thin layer of the PS were obtained by anodic etching. The monocrystalline silicon plates of p-type and electro- lytes of hydro-fluoric acid were used. The thickness of the PS-layers was about 100–400 A0. Human Mb extracted from the heart muscle and specific mouse mono- clonal Ab were used. These components were prepared in concentrations of

Biosensors for Biochemical Diagnostics of Diseases 467 20 μg/ml in 10 mM PhB, pH 7.3. For the quantitative analysis the changes of the I-V parameters were investigated (ΔIі/I) as a function of voltage and thickness of Ni and PS-layer (where: I is current through structure with- out drawing substrate and ΔIі is the difference between currents through structure before and after drawing appropriate substrate. The index i meets following: bulk structure (0) and after the deposition of Mb (1), its specific monoclonal Ab (2), consequent deposition of Mb and it’s Ab (3), mixture of Mb and Ab (4) on the surface). After the deposition of Mb or its specific monoclonal Ab on the super- thin Ni film the changes in the I-V characteristics of the SBS are observed. The response of these structures varies more essentially when the specific immune complex was on the surface (after the consecutive deposition of Ag and then Ab). For all investigated types of biological molecules and their specific complexes the characteristic optimum thicknesses of Ni film and PS-layer that give the largest changes of the I-V characteristics are equal to 200 – 300 A0 and about 200 A0 respectively. The morphology of Ni or PS surface influences on the structure of active centers of biological molecules that leads to the changes in electrometric characteristics of the SBS. Also it is possible to assume that after the deposition of protein com- pounds conductivity through pores or some deepening’s of Ni film, filled with a biological substance, dominates. The current research demonstrates a possibility of immediate detection of a specific immune complex in electrometric way. The data obtained in this research open a perspective to use the electrometric approach as a basis for the creation of new types of immune sensors. 13.2.6 Immune Biosensor Based on the Surface Plasmon Resonance (SPR) Direct control of biospecific interactions is very important for solving of a number of tasks in the field practical medicine in the respect of application of the new methods for diagnostics. The optical technique, in particular based on the SPR phenomenon, is the most promising. Kretschmann in 1973 has proposed[26] the attenuated-total-reflectance method which is widely used now. In 1983 it was demonstrated for biosensing [27]. From that time, the SPR method gives possibility to monitor the association and dissociation of a wide range of biomolecular complexes in real time with- out the use of labeled molecules. In this technique light is totally reflected at a glass-metal film interface and the reflectance is monitored as a function of incidence angle (resonance curve). At a certain angle, a minimum in the intensity of reflected light is observed. This indicates that the coupling of

468 Advanced Biomaterials and Biodevices energy occurs between the incident light and the surface plasmon waves. The position of the resonance angle is sensitive to the changes in the refrac- tive index and/or thickness of the layer in the vicinity of metal surface. When the resonance angle varies in time, we obtain kinetic curve. The usage of gold layer as a transducer in the SPR biosensors is very perspective since it allows obtaining stable and reproducible signal. Acceptable performance of the sensors requires, even for gold films, a pro- tective coverage to improve the stability of transducer during immersion in aqueous solutions and smoother surface. Some proteins are difficult to attach directly to a metal surface, some of them may denaturate at the direct adsorbing on such surface [28]. In our investigations we realized a lot of different methods for biological component immobilization on the transducer surface [29–31]. 13.2.6.1 Estimation of Anti-insulin Ab in Serum Blood [32–34] Today we know that the presence of Ab against any structures of own organism shows about the development of autoimmune disorders. That is why, the problem of the identification of this class of compounds and their quantitative analysis has important scientific and practical signifi- cance, because it allows, on the one hand, studying the mechanism of various pathological forms of pancreatic cancer, knowledge of which will promote individualization and improve insulin therapy from other one. This result should be provided as soon as possible even with very high sensitivity. The presence of Ab to insulin may interfere with the determina- tion of its level in the blood and affect the outcome of glycemic control in patients with diabetes. In addition, high levels of Ab against insulin may contribute to the immune resistance in these patients, requiring the need for dynamic control of the concentration of this type of Ab in patients with diabetes who receive various insulin preparations. Especially it should be noted that in most cases there is a need to use express (sometimes even result should be obtained immediately), highly sensitive and specific meth- ods strictly. Just the last two requirements are met by classical methods of modern immune-chemical analysis, such as radio-immune (RIA) and immune-enzyme (the ELISA) methods. The latter is used for the detection of Ab to glutamic acid decarboxylase [35]. Its sensitivity reached 20 ng/ml. However, in both cases the expressivity of the immune-chemical methods is insufficient. In full it could be really provided only with the immune biosensors. Early [36] for the detection of auto-Ab to insulin the immune biosensor based on ellipsometry was developed which has sensitivity in the range of 10 ng/ml to 100 μg/ml.

Biosensors for Biochemical Diagnostics of Diseases 469 Application of the biosensor based on the principles of the SPR [37, 38] will enable professionals engaged in the diabetology, implement a radically new strategy for diagnosing and studying the pathogenic mecha- nisms of desease, which is as “the number one problem” in endocrinology. The principle of biosensor work was described in detail early [32- 37]. Transducer was as glass plate with the gold (20 nm) deposited on the previously formed layer of chromium (3 nm). It was connected with the prism of the measuring device through polyphenyl ether with the refrac- tive index in 1.6. IgG and insulin dissolved in 1 mM PBS (pH 8,2) at the concentration of 1 μg/ml were used as specific Ag. In each experiment it was registered resonant angle at the successively introduction in the mea- suring cell: distilled water, specific Ag, solution of BSA in PBS at the con- centration of 1 mg/mL and antiserum with the dilution from 1:10000 to 1:200. The time of the exposition was 20 min and after each measuring the cell was washed by distilled water. It was shown [36] that modifications metal surface by the adsorption of the Ag was stable over time and was not destroyed by washing mea- suring cell with PBS. Immobilization Ag was accompanied by changes in the resonant angle within 3200–3500 arc sec. In the case of polyelectro- lytes amount of Ag adsorbed on the surface was slightly larger as well as biosensor was more stable and reproducible than one with bulk surface. Application dodecanthiol also helps to stabilize the immobilized layer of protein and increases their density on the surface. Number of physically adsorbed biological molecules is limited by surface area of the transducer, which is amenable to optical registration. This limitation can be prevented by the location of biological material in three-dimensional space. Of course, the amount of immobilized material can be increased, and thus it is possible to further increase the sensitivity of the biosensor. To achieve of such situation it was therefore proposed a different scheme using polyca- tions and polyanions [30, 31]. Solution of insulin at the concentration of 0.5 μg/mL in 1 mM TB (pH 8.2) was introduced into a measuring cell for 20 min at the room temperature. Then the cell was washed by PBS and filled successively polyethylenamine chloride (PAA) – polystyrene sul- phate hydrochloride (PSS) – PAA at a concentration of 1 μg /mL in water. Cell was washed by water and treated by insulin during 20 min and washed again. At last it was filled by 0.5% solution of GA for 15 min and polyelec- trolytes were removed by 50 mM buffer (pH 4.0 and 8.0) to form two-layer sensitive surface. It was established that the formation of a double layer of insulin on the transducer surface increased the sensitivity of the immune biosensor to specific Ab. This is particularly important in the study of the early stages

470 Advanced Biomaterials and Biodevices of the disease when specific Ab titre is low. However, a complication of this process is not appropriate for practice, because the main advantage of biosensor analysis belongs to its expressivity. Thus, it was analysed the effectiveness of selective immobilization of biological material on the sur- face previously covered with thiols, PAA, PSS, separately. It was stated that the use of PAA for the modification of the transducer surface is the most appropriate, affordable and simple. It was found [32, 33] too that the most acceptable is the algorithm of analysis, namely when insulin immobilized on the surface of the optical transducer directly interacts with specific Ab. To find the optimal concentrations of the components for the biosensor application it was prepared samples of insulin and Ab from different sources (from two different firms) and recombinant insulin in 0.01 M PhBS (pH 7.4). It was stated that the optimal concentration of recombinant insulin should be considered about 20 μg/ml and for pig insulin there is necessary dilution in 10 times. With the used insulin samples it was constructed the needed calibration curves. To avoid the non-specific interaction of biomoleculs the dilution of serum should be not less than 1:100. The samples from ill persons were divided by the standard ELISA-method according to levels of antibod- ies against insulin (the maximum and minimum). As it can see from the data presented in Figure 13.6 there was a very high correlation biosensor response with predefined initial levels of Ab against insulin in samples. So, if all procedures needed for the analysis, including pre-processing transducer surface and insulin immobilization will perform simultane- ously with the sample analysis it is necessary to spend more than 30–40 min. But this time is much less that needed to perform the traditional ELISA-method (4–5 h). Moreover, in case of biosensor method there is a possibility to fulfil preliminary procedures as the previous surface trans- ducer treatment and immobilization of insulin and then pure time of anal- ysis will take only 5–10 min. 100 50 0 123 Figure 13.6 Response of the immune biosensor at the analysis of antiserums (dissolved in 1:100) from the immunized pig (1), patients with maximal (2) and minimal (3) content of antibodies. Abscissa – changes of reflectance angle, min.

Biosensors for Biochemical Diagnostics of Diseases 471 13.2.6.2 Determination of Mb Concentration [37] The quantity of Mb in biological liquids (blood, urine) can tell us about series of human illnesses, for example: ischemia, infarction, some meta- bolic disturbance. The technical aspects were the same as it was described above at the determination of anti-insulin Ab. But in this case a freshly prepared gold surfaces were modified with the thioglycolic acid anilid - C8H9ONS (short molecule) and the dodecanethiol - HS(CH2)11CH3 (long chain molecule), which differ in length of the hydrocarbon chains and nature of the tail groups. Chemisorption was performed by 15 h immersion of gold films in 1mM ethanol solution at room temperature, followed by thorough rinsing with absolute ethanol and drying in a flow of clean air. The overall time of incubation with the biocomponent was no more than 5 min. It was investigated the binding of the specific monoclonal Ab with human, rabbit and badger Mb’s. On the surface of transducer the mono- clonal Ab (with the concentration of 200 μg/ml) were immobilized by the physical sorption from a solution in a 50 mM PBS, pH 7,4. Unbounded Ab was displaced by the PBS and then the corresponding Mb with the different concentration was added. It was shown that the cross reaction of the inter- action of non-specific Ab with the any Mb was negligible. The shift of the resonant angle was already observed at the injection of 1 ng/ml of appro- priate Mb to the immobilized monoclonal Ab. The sensitivity of analysis was in the range from1ng/ml to 10 μg/ml. The formed immune complex did not destroy by PBS washing. Total dissociation of specific interaction was achieved after the injection of the buffer with pH 2,2 into system. After that we observed only presence of adsorbed Mb on the surface. The storage of the preliminary prepared transducers at room temperature for about 5 day did not deteriorate its sensitivity. The most appropriate results were obtained at the thiol-modified gold surface application. Thus, there is all reasons for the conclusion that this biosensor may find application as the suitable method of diagnostics of various illnesses based on the estimation of Mb content in biological liquids. 13.3 Electrochemical Enzymatic Biosensors Based on the Ion-sensitive Field Fffect Transistors (ISFETs) Many types of enzyme biosensors based on the ISFETs were cre- ated after the first reported work in 1980 [39]. In the comparison with the other types of biosensors, the ISFET based ones have certain

472 Advanced Biomaterials and Biodevices well-known advantages: miniaturisation, high sensitivity, low cost and multi-detection potential (this is especially important for the creation of multi-biosensors). Such biosensors were adapted for different purposes including glucose and urea determinations [40–44], with the idea of fur- ther use in clinical analyses. It can be very important to have such bio- sensors in many clinical cases dealing with diabetes or kidney and liver diseases, when fast and very sensitive glucose or urea measurement is essential. Unfortunately, the silicon nitrate ISFETs shown not higher sta- bility and now we propose to change them by that based on the cerium oxide. The last were applied already by us for the immune biosensors at the control of S. thyphymurium level [45]. Since the principle of the application, obtained results are similar in case of work with both types of ISFETs we will pay attention now such type of biosensors intended for glucose and urea control. 13.3.1 Analysis of the Urea Level in Blood [46] The immobilized urease stipulates pH changes around the gate surface according the reaction: (NH2)2CO + 2H20 + H+ urease 2NH+4 + HCO3- Sensor chips with two ISFETs were fabricated by N-channel LOCOS- technology. The dimensions of a p-silicon wafer were 3 mm wide, 10 mm long and 0.3 mm thick. Sensor sensitivity was linear from pH 3 to pH 10 with a slope of 40–45 mV/pH unit. The procedures for making ISFETs and their properties were reported previously [39–43]. Urease was immo- bilized from solutions with BSA at a concentration of 50 mg ml-I in in 20 mM PBS, pH 7.4 and 2.5% glycerol mixed in equal volumes. A drop of this mixture (0.1 μl) was deposited on the sensitive area of the sensor. A similar membrane prepared only with BSA in buffer solution was placed onto the reference ISFETs. To complete a polymerization of the biomem- brane the sensor chips were transferred to a saturated vapor of GA for half an hour. All the experiments were performed at 23 °C in a 2.5 ml measuring cell under constant stirring. Sensor was immersed in the work- ing buffer for 0.5–1 h before use. Between the experiments sensors were stored at 4 °C in 20 mM phosphate buffer pH 7.4 with 1 mM ethylenedi- aminetetraacetic acid (EDTA). The differential measurement method was used to eliminate the non-specific influence of the external conditions (temperature, light and pH fluctuations) on the biosensor signal. A typical time-response curve of the urea biosensor had an exponential character. The maximum signal was achieved after 1–3 min. This time depends on the membrane thickness. The output of the urease sensor in 1 mM urea solution had a maximum at pH 7.0, which is within the range for optimal

Biosensors for Biochemical Diagnostics of Diseases 473 urease activity (pH 7.0–7.5) and close to the human blood (pH 7.4). This makes possible the measurement of blood samples without further adjust- ment of pH value. The biosensor response for urea also depends on the ambient buffer capacity, the signal being sharply reduced by an increase in the buffer concentration from 1 to 10 mM. However, the linear part of the calibration curve is extended for higher buffer concentrations, making readings for 0.25–1.50 mM urea possible. Because the buffer capacity of blood is relatively high (due to the presence of proteins and buffer salts) reductions of biosensor output can occur during the addition of blood samples to the measuring cell. To avoid the effects of blood buffer capacity, the concentration of working buffer in the measurement cell should there- fore be not less than 5 mM. The main salt component of blood is sodium chloride, with a concentration of 150 mM. This conditions strongly affect on the efficiency of the urea biosensor. The biosensor output falls by 50% in the presence of 200 mM NaCI and stays constant during further addi- tions up to 500 mM NaCI. Thus to eliminate the influence of the salt concentration of blood samples, a working buffer should contain at least 200 mM NaCI. Using these selected parameters from model solutions, the urea biosensor has been tested on blood samples from laboratory animals. In such experiments a 25-fold dilution of serum samples has been found to be optimal for obtaining a reliable response. The presence of serum in the working buffer changes the slope of the calibration curve, which is probably due to the buffer capacity of blood proteins. It was created the calibration curve to estimate the urea concentration in animal blood samples and the results were compared with those obtained by traditional chemical way. We have obtained a very high correlation for both methods. Such biosensor can be used 20 times without a decrease of the response value. Deviations between individual response values are about 10%. After 130 measurements a 30% decrease in response value was observed. Thus, this urea biosensor seems to be suitable for further testing in med- ical practice and especially if the silicon nitrate in ISFETs will be changed on cerium oxide as it was mentioned above. 13.3.2 Determination of the Glucose Level in Blood [47] Now there are a several dozen papers dealing with enzymatic ISFETs based biosensors. They are based on the following reaction catalysed by β-glucose oxidase (GOD) immobilized on the gate dielectric of the ISFETs: β-D- glucose + O2 β-glucose oxidase D-glucono-δ-lactone + H2O2. Subsequent hydrolysis of gluconolactone giving gluconic acid (pK~3.8) changes the solution pH value near the sensor surface.

474 Advanced Biomaterials and Biodevices At least four major obstacles are in the way of the applications of glucose ISFETs based biosensor: (a) the dramatic influence of the buffer capac- ity of a sample on the sensor response; (b) the non-linear dependence of the enzyme kinetics and buffer capacity of a pH sample; (c) influence on the response ionic strength; (d) co-substrate limitation of the enzymatic reaction. Regarding the last point, the glucose concentration in human blood is normally about 5 mM, reaching 20 mM and more for diabetics. However, the concentration of oxygen dissolved in a sample, as a rule, does not exceed 0.5 mM. So because of the unfavorable ratio of glucose and oxygen in real blood, the dynamic range of the biosensor is usually limited by oxygen. To avoid this situation with the “oxygen deficit” the analyzed samples were preliminary diluted. Sensor design, creation of biomembrane and measurements were made similar to that as it was described above in case of urea analysis. In addi- tional to biomembrane was prepared with the help of polyacrylamide gel. Both membranes showed good results at the biosensor application. The substrate sensitivity of this biosensor reached a maximum of ~10–5 M of glucose in unbuffered medium and depends dramatically on the buffer capacity of the sample. This is due to the presence in the sample of a mobile pH-buffer that augments the flux of H+ ions produced by the enzymatic reaction out of the membrane by means of the ‘carrier-mediated’ transport mechanism (“facilitated diffusion”) [48]. The dependence of the biosensor response on the ionic strength of the sample is characterized by the next data. The curve levels off at [NaCl] >150 mM and does not change when using MgCl2, or KC1 instead of NaCl. This dependence tends to linearize when plotted against the logarithm of the ionic strength, thus implying the mainly electrostatic nature of the effect. It is also shown that there is no substantial effect of the protein content of a sample on the biosensor response. Since the solution buffer capacity is strongly influenced by proteins, it follows that only those components of a buffer that are sufficiently small and mobile to penetrate inside the bio- membrane can affect the sensor response. The obtained results allowed us to recommend the following protocol for the glucose determination in blood by the ISFETs-based biosensor: (I) to overcome the problem of ‘oxygen deficit’ and to adjust the glucose con- centration in a sample to the sensor quasi-linear operating range (0.1–2.0 mM), blood should be diluted in 10–20 times: (2) since the concentration of buffer in blood does not exceed 30 mM (mainly bicarbonates) there- fore diluting the sample with, e.g., 2–5 mM Na-phospste buffer (Na-PhB) (pH 7.3–7.5) should be done (in this case the buffer capacity of the result- ing solution will be mainly determined by the diluent); (3) to stabilize the

Biosensors for Biochemical Diagnostics of Diseases 475 ionic strength of the sample, NaCl should be added to the diluent up to 150 mM concentration (this procedure may be optional when using 5 mM as a diluent, since its ionic strength is sufficient to stabilize the sensor response within the 10% relative accuracy usually required); (4) the almost invariable sensor response in the pH range 7.0–8.0 makes it unnecessary to adjust the sample pH to a certain value or to calibrate the sensor at a predefined pH value in this range. The biosensor was examined according to the next protocol at the blood analysis: (a) the drawing baseline of biosensor immersed in diluting buffer (2 mM of PBS, pH 7.4) then the registration of its response for 1 mM glu- cose concentration, further after rinsing, 0.1 volume of undiluted rat blood was added and afterwards the sensor response for 1 mM glucose should be examined again; (b) in another test after adding 0.1 volume of rat blood, a solution with native GOD was added which eliminated the glucose from the reaction volume and the base level signal of the sensor was restored; (c) the last test included addition of 0.1 volume of rat blood already treated with GOD, no signal was obtained in this case. Subsequent addition of 1 mM glucose results in a correct response value. The biosensors did not exhibit any essential loss of activity after 100 assays. It showed less than 10% loss of activity after two month storage (in 0.5 mM PhB, pH 6.5, at +4 0C). The lifetime of the sensors can be signifi- cantly increased at the storage in a dry state and at +4 0C. In general, it is possible to conclude that the presented system is suf- ficiently versatile to meet any requirement concerning accuracy, response time, stability, calibration and reproducibility of biosensor. There is possi- bility of applying the developed biosensor to blood glucose assay as incor- porated in a flow-through system or as a disposable probe. The very small volume of blood (up to 10 μl) and short time (less than a few seconds) needed for one glucose test make such biosensor very attractive for bio- medical applications. 13.4 Multi-parametrical Biosensors [49–51] To provide creation of the analytical devices which will correspond to full complex of practice demands including selectivity, sensitivity, stabil- ity in general, repetition of results, needed time of functioning and oth- ers it was developed microelectronic enzymatic biosensor which gives possibility to simultaneously register a number of parameters, namely, the determination of glucose, sucrose and lactose in case of control of technological process at the sucrose production or the estimation of

476 Advanced Biomaterials and Biodevices glucose, insulin and anti-insulin Ab level at the control of the diabetic and autoimmune state. Amperometrical channel includes whole line of planar microelectrodes with help of which current was registered. As a rule the average value of three microelectrodes was preceded. Thermo metrical channel is presented by thin layer sensors which work in differential regime. Output signals of the integral biosensor are processed by the special electronic block. Thin layer temperature sensors were prepared by two ways: on the basis of plati- num or oxide cerium. These sensors together with the platinum working electrodes and ref- erence electrode (AgCl) for the amperometrical measuring were situated on the common solid phase. Three working electrodes were taken for the increasing of accuracy of measurements. Thermo metrical channel con- tents also working and reference thermistors which have weaving form. Enzymes as well as Ag or Ab were immobilized on the transducer surface after its preliminary treatment by PAA or PSS. The integral biosensor contents simultaneously two separate channels: amperometrical and thermo metrical is presented in Figure 13.7. It was demonstrated that thermometry allows providing control of concentrations in frame of 0.002–25; 0.005–100; 0.1–10 mM and sensitiv- ity about 2.35; 0.45 and 0.85 rel. units/mM for glucose, sucrose and lactate Protective layer (B) (AI O ) 23 Pt Ag AgCI Pt 0.1 Sensitive layer (C) 15 mm 0.05 0.1 0.05 0.05 Pt 4.6 mm (A) 1 23 4 5 6 (D) Figure 13.7 Scheme of the microelectronic chip (with concrete materials and structure dimensions (A), overall view of two parts of the multi-parametrical biosensor: (B) – electronic unit for the signal registration and (C) – multi-channel block with the microelectronic registration of the ampero- and thermo metrical signals. (D) – the section of the biosensor: 1-quartz basis; 2 – layer of AgCl; 3 – silver electrode; 4 – measured platinum electrodes; 5 platinum thermo resistance; 6 – biochemical structure (additional explanations in the text).

Biosensors for Biochemical Diagnostics of Diseases 477 correspondingly. From other side amperometry was characterized by values of controlled concentrations in frame of 0.5–50; 0.1–80; 0.1–200 mM and sensitivity on the level of 4.3; 3.3; 3.3 nA/mM for above mentioned sub- stances, respectively. In general it is necessary to mention that these values correspond to the practice demand in respect of the sensitivity from one side and from other one the obtained values have been confirmed by special bio- chemical experiments. It is very important to underline that the total time of analysis is very short and in additional to there is possibility to obtain irrespective data about level of some parameter by different approaches. The sensitivity of the determination of insulin and anti-insulin Ab in serum blood of patients with diabetics is highest at the application of the thermo metrical channel of the multi-parametrical biosensor among other ones. It was chosen the optimal insulin concentration for the immobilization. Then the calibration curves were constructed for both channels (Figure 13.8) and at last, it was examined the efficiencies of work both channels at the analysis real samples preliminary tested by the ELISA-method (Table 13.1). The obtained results on the beginning of our investigation allow us already to conclude that such types of the multi-parametrical and 10 Sensor response, 200 5 mK 150 I, nA 100 0 0,5 1 1,5 2 0,1 0,25 0,4 0,55 0,7 0,85 1 0 C, mkg/mL 50 c, mkg 0 (A) (B) Figure 13.8 “Direct” detection of anti-insulin Ab in the serum blood by the thermo metrical channel (A) and by the amperometricl channel (B) of the multi-parametrical biosensor. The concentration of the immobilized insulin was 10μg/mL and second Ab were labeled by the HRP. Table 13.1 The work efficiency of the multi-parametrical biosen- sor at the discovering results preliminary stated by the ELISA- method (10 samples). Method of analysis % of discovered results ELISA-method 100 Amperometrical channel 94 Thermo metrical channel 85

478 Advanced Biomaterials and Biodevices multi-functional biosensor will be able to provide all practice demands in the respect of the screening of several objects on the number of biochemical indexes. In medical diagnostics it may be used for the simultaneous detec- tion of the glucose concentration and the anti-insulin Ab in patients with diabetes as well as for the registration of others biochemical parameters. 13.5 Modeling Selective Sites and their Application in the Sensory Technology The development of biosensors is limited by the accessibility of the suitable biological material and its insufficient stability in solution under the influ- ence of high temperatures and some chemical reagents which decrease the sensitivity and selectivity of such devices. To avoid these disadvantages it was used two different approaches: one of them connected with the cre- ation of the selective sides by the special chemical way [52, 53]. It gives possibility for formation of the special places on the solid phase which content the optimal environment in relation to steric and other interac- tions which can provide binding analyzed substances. Second one is based on the application of chemically structures which have some selectivity to analyzed substances. In both cases there is necessary to use a lot of such semi-selective sites and to apply a special computer program for discrimi- nation of non-selective bindings and revealing selective ones. 13.5.1 Template Sensor: Principle of Creation and Characteristics of Work and Determination of Some Biochemical Substances [52] For the realization of this purpose the silicon plates were treated with hot steam during 1 h. Then the plates were placed in a 5 mM solution of the templates for 12 h. After removing the excess solution the plates were dried at 120 0C for 3 h and placed in a vacuum flask with 5 ml of frozen Me3SiCl and kept at room temperature for 12 h. Finally, the templates were washed from the plates with water and ethanol. As a control the samples modified by silanes but without templates were used (Figure 13.9). The glass electrodes were modified with Me3SiCl as described above and by ODSiCl3, as reported previously [53] in the presence of cholesterol and without it (control samples). The indium tin oxide (ITO) electrodes were modified by Me3SiCl in the presence of phenylalanine. The sorption of [P32] ATP or amino acids was carried out on modified plates in the presence or absence of template for 30 min. Then the plates were washed very carefully

Biosensors for Biochemical Diagnostics of Diseases 479 (A) (B) Figure 13.9 Scheme of measuring arramgement (A), where: 1 – injector, 2 – probe, 3 – Pt electrode, 4 – Ag/AgCl electrode, 5 –analyzed surface, 6 – contact to analyzed surface and surface modification (B). with water, dried and their radioactivity was measured. The experimental set-up used to investigate the conductance of the layer on the electrode sur- face consisted of a potentiostat, a flow injection electrochemical cell and an X-Y recorder. The teflon probe with injector and saturated Ag/AgCl reference electrode was used for realization of the flow injection electrochemical cell. This construction allows the liquid analyte around the electrode surface to be changed rapidly. The probe working surface is about 0.75 cm2. Measurements were performed with tris-HCl buffer solution as the background electrolyte. Solutions with concentrations of 100 mg/l Phe were achieved by addition of the corresponding substances to the background electrolyte. The working electrode was controlled at +0.9 V (respectively to Ag/AgCl RE). The efficacy of this approach was confirmed by the analysis of the sili- con plates treated with silane in the presence of adenosine (A), AMP, ATP and TTP on their sorption ability for P32ATP. The “ATP-modified” plates had a high selectivity for their substrate-template. The sorbtion of leucine, methionine and phenylalanine is largest for the surfaces modified with the corresponding amino acid template. The difference in specific sorption of phenylalanine and leucine by the sample modified with leucine is not very large. Also, the specific sorp- tion of phenylalanine by Phe- and Leu-modified samples is very similar. It cannot be related to the homologous structure of the molecules such as phenylalanine and leucine. Probably, the sorption of molecules on the tem- plate sites is an integral value governed by molecular shape and structure and by the interaction between the functional groups of site structure and sorbet molecules. The sorption of cholesterol on electrodes modified with silane in the presence of cholesterol leads to the surface monomolecular layer being

480 Advanced Biomaterials and Biodevices filled and an increase of pH on titration of a 50 mM water-methanol (1:l) solution of Na2HPO4, by dinitrophenylacetic acid. For electrodes modified by ODSiCl3, the difference in Ph was 0.15, for Me3SiCl-modified electrodes it was 0.1 and for non-modified electrodes it was 0.05. (The concentration of cholesterol was the same in all cases, 2 mM). The electrodes modified by silane without the presence of cholesterol, do not show any response to cholesterol. In addition there is no response of cholesterol-modified elec- trodes to phenol and aromatic amino acids, but there is the same response to cholesterol-homologous compounds such as litocholeic acid and deoxy- choleic acid as there is for cholesterol. It was shown that the hydropho- bic layer is destroyed after using the modified electrode in such butlers as NaH2PO4, NaHCO3 and CH3COONa. It leads to a decrease of pH. Perhaps, a better result could be achieved when the monomolecular film of silane is cross-linked in the presence of a template. We have a preliminary result, the template-treated field-effect transis- tor gives a response for cholesterol at a concentration of 10 mol/l. Some “Phe-modified” electrodes were used for an investigation of the receptor structure. In the case of undeveloped IT0 and complete silane modified IT0 an injection of amino acids caused an increase of anodic current. All current responses were reversible. As expected modification of the IT0 sur- faces by organosilane layers caused a decrease in conductance. In the case of template-modified IT0 surfaces an injection of amino acids, except for Phe, increases the anodic current as for previous samples. In contrast to this an injection of Phe caused a decrease in the anodic current. This cur- rent change is not reversible. It shows the specific binding of Phe to selec- tive sites on the template-modified surface. The current is reversed after an injection of ethanol during 5–6 min. In summary, a simple method of creating a selective SiO, surface for the specitk sorption of small-molecular organic compounds is proposed. The method is based on a preliminary sorption of the template, treatment of the surface with silane and washing the formed structure. This method is useful for immobilization of unstable biological materials. We hope the present procedure will be useful for creating new sensor devices for use in medicine, biotechnology and environmental pollution monitoring. 13.5.2 Artificial Selective Sites in the Sensors Intended for the Control of Some Biochemical Indexes [54] One of main among others approaches for the perspective development of the new generation of the instrumental analytical devices is the replacing biological sensitive biological materials by the artificial structures on the

Biosensors for Biochemical Diagnostics of Diseases 481 Figure 13.10 Calyx[4]arenas (bellow) and possible type of binding of molecules on the surface covered by gold and modified by dodecanthiol (above). basis of calyx[4]arenas [54]. Certainly, as it was mentioned above, in this case the absolute selectivity will be absent but there is possible to obtain specific signals by the application of the analysis of the number of the dif- ferent step of the non-specific ones by the help of the special computer program. The overall view of the calyx[4]arenas and their immobilization of the gold surface of the transducer through dodecanthiol is shown in Figure 13.10. It was analyzed the binding constants of the calyx[4]arenas with some substances and was shown that they have strong differences which have possibility to discriminate the formed complexes. Of course, it is possible to synthesis a lot of variants of the calyx[4]arenas which will have much more differences in the affinity to the separate substances. For the realiza- tion of the proposed principle there is necessary to use a principle of multi- parametrical and multi-functional analytical devices. 13.6 Conclusion Serological investigations have of great practical importance for the pri- mary screening of a number of biochemical indexes at the different stages of the diagnostics of pathological states of organism to reveal of a nature of

482 Advanced Biomaterials and Biodevices disease, for example: some defects in lung, heart, liver, kidney e.a. The prog- ress in the serological methods are closely linked to the development of the enzymatic and immune biosenors that allow fulfilling an analysis in such mode as it requires practice, namely, sensitive, specific, fast and cheap. An important advantage both mentioned types of biosensors are that analysis can be done in real time and in the field conditions, at the patient bad or at home directly. Today a lot of types of biosensors are developed by the real- ization of a different physical and chemical principles for the registration of the specific generated signals. There a great significance plays the choosing effective ways of the modification of biosensor surface, optimization of the conditions for immobilization of biological material and choosing the most available protocol of analysis. At the achievement of the better parameters for the above mentioned positions we can obtain the sensitivity analysis which exceeds that which may be provided by the traditionally used meth- ods. Moreover biosensors may give the advantages, as a rule, before the tra- ditional methods in time which should spend for the analysis. The duration of analysis using biosensors is in frame of 40 min, including time for the immobilization of the sensitive elements on the transducer surface, block- ing free binding sites and rinsing the measuring cell. The some modification of protocol of the analysis fulfillment can give possibility to short this time up to 5–10 min. Device based on the biosensor principle can be as portable. In generally it was demonstrated that the practical realization of the bio- sensoric principles (using the optical and electrochemical transducers at the creation of the instrumental analytical devices) may provide simple, cheapness, express, selective and sensitive analysis. The possibility to apply artificial substances, in particular, using calyx[4]arenas, some template structures as selective sites and the development of multi-parametrical and multi-functional biosensors are considered as further perspective of the creation of the instrumental analytical devices. The very important role belongs to the use of nano-structured materials, nano-particles and nano- tubes in form of transducers of new generation. References 1. N.F. Starodub, P.Ya. Arenkov, A.E. Rachkov and, V.A. Berezin, Sensors and Actuators B, Vol. 7, pp. 371–375, 1992. 2. E. Engvall and P. Perlmann, Immunochemistry, Vol. 8, pp.871–874, 1971. 3. M.J. Martin, V.A.B.D. Wickramasinghe, T.P. Newson and J.A. Crowe, Fibre- optics and optical sensors in medicine, Med. Biol.Eng.Comp, Vol. 25, pp. 597–604, 1987.

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14 Nanoparticles: Scope in Drug Delivery Megha Tanwar, Jaishree Meena and Laxman S. Meena* CSIR-Institute of Genomics and Integrative Biology, Delhi University Campus, Delhi, India Abstract Nanotechnology is the most expeditiously emerging technology in the field of therapeutics as drug delivery systems. It is the manipulation of matter on atomic and molecular scale. It works on devices, structures and materials with at least one dimension sized from 1 to 100 nanometers. Nanoparticle based technology has played a significant role in treatment and prevention of tuberculosis, cancer, etc. Recently a variety of nanocarriers has been evaluated as potential drug delivery systems for various administration routes. Liposomes and microspheres have been developed for sustained delivery of anti-TB drugs. Several anti-cancer drugs includ- ing paclitaxel, doxorubicin, 5-fluorouracil and dexamethasone have been success- fully formulated using nanomaterial. Targeting of drugs to certain physiological sites using nanoparticles has emerged a promising tool in treatment of tuberculosis with improved drug bioavailability and reduction of dosing frequency. Nanotechnology based targeting of drugs may improve the therapeutic success by limiting the adverse drug effects and resulting in more patient’s compliance and attaining high adherence level. Nanotechnology holds merit as a drug carrier because of its high carrier capac- ity, feasibility of incorporation of both hydrophilic and hydrophobic substances, high stability, and also because of feasibility of variable routes of administration like oral application and inhalation. The nanoparticle based drug delivery systems are advantageous over other modes of drug administration as nanoparticles deliver the drug more efficiently and with reduced side effects as well. Most recently, various vesicular systems have been developed such as niosomes which can be explored for achieving maximum effective concentration of the delivered drug. They can be uti- lized for research on various drugs and may turn out to be most promising mode of drug delivery in treatment of various deadly diseases. Keywords: Nanoparticles, targeted drug delivery, nanomedicine *Corresponding author: [email protected] Ashutosh Tiwari and Anis N. Nordin (eds.) Advanced Biomaterials and Biodevices, (487–522) 2014 © Scrivener Publishing LLC 487

488 Advanced Biomaterials and Biodevices 14.1 Introduction The development of nanoparticles for drug delivery began in the1960s [1]. Recent developments in multifunctional nanoparticles has offered a great potential for targeted delivery of drugs for treatment of various types of diseases. Nanoparticles are basically solid colloidal particles ranging in size from 1 to 1000nm (μm). Nanoparticles are made from biocompat- ible and biodegradable materials such as polymers, either natural (e.g. gelatin, albumin) or synthetic (e.g. polylactide, polyalkylcyanoacrylates) or solid lipid nanoparticles [2]. The drug loaded in nanoparticles is usu- ally released from the matrix by diffusion, swelling, erosion or degrada- tion. The reason, why nanoparticles used for medical purposes is large surface to mass ratio which is much larger than other particles. Large surface area provides ability to bind, adsorb and carry other compounds such as drugs, proteins, probes. They consist of macromolecular materials in which active agent (drug or biologically active material) is dissolved, entrapped, encapsulated or to which active agent adsorbed or attached [3]. Unfortunately the conventional therapeutic strategies require unnecessar- ily high systemic administration due to non-specific biodistribution and rapid metabolism of free drug molecules before reaching their targeted sites. Therefore nowadays nanotechnology based targeting improved the therapeutic success by limiting adverse drug effects and patient requires less frequent administration regimes, which ultimate results in more Oral Injectable Routes of Implantable administration of various drug delivery systems Inhalable Figure 14.1 Various Routes of Drug Delivery Systems.

Nanoparticles: Scope in Drug Delivery 489 patients compliance and thus attaining higher adherence levels [4]. There are many advantages of nanoparticles as drug carriers including high stability, high carrier capacity, and feasibility of incorporation of both hydrophilic and hydrophobic substances, feasibility of variable routes of administration including oral administration and inhalation and reduc- tion in dosing frequency [5]. Mainly uptake of nanoparticles occurs by three mechanisms (1) transcytosis (2) intracellular uptake and transport through the epithelial cells lining the intestinal, mucosa, (3) uptake via peyer’s patches [6, 7]. There are various routes of administration of drug delivery systems as described above in Figure 14.1, which are advantageous in one or another way. 1. Oral drug delivery-Most acceptable route for drug delivery, substantial reduction in dosing frequency, economic and ease of preparation [8]. 2. Implantable drug delivery-High drug bioavailability and least dosing frequency [8]. 3. Injectable drug delivery-It also has high drug bioavailabil- ity and least dosing frequency but painful procedure [8]. 4. Inhalable drug delivery-Direct drug delivery to the target site, reduction in dosing frequency and improved bioavail- ability [8]. 14.2 Different Forms of Nanoparticles as Drug Delivery Nanoparticles are used in various forms for drug delivery such as nano- spheres, nanoemulsions, nanocapsules, solid lipid nanoparticles, nanosus- pensions, polymeric nanoparticles, liposomes and micelles, dendrimers and niosomes. Monolithic nanoparticles (nanospheres) are those in which drug is adsorbed, dissolved or dispersed throughout the matrix and nano- capsules are those in which drug is confined to an aqueous or oily core sur- rounded by shell like wall [9]. Nanoemulsions referred as miniemulsions or sub-microemulsions by dispersing mainly oil in water. Thermodynamically stable nanoemulsion (mean size 80.9 nm) of ramipril was developed for oral administration. In vitro studies showed that drug release till 24h from nanoemulsion band was much more important as compared to marketed capsule formation and drug suspension [4]. The relative bioavailability of ramipril nanoemulsion to that of conventional capsule was 229.62% and to

490 Advanced Biomaterials and Biodevices Liposomal Free drug drugs Reduce Stratum irritation Corneum Epidermis Enhance drug permeation Dermis Prolong Blood resistance time supply Reduce systemic toxicity Figure 14.2 Liposomal drug and free drug delivery compared. that of drug suspension was 539.49% which suggested the use of developed ramipril nanoemulsion for paediatric and geriatric patients [10]. Solid lipid nanoparticles are lipid-based submicron colloidal carriers. They were initially designed as pharmaceutical an alternative to liposomes and emulsions. In case of solid lipid nanoparticles, the drug is entrapped in solid lipid matrix to produce lipid nanoparticles of size range 50–100nm by using hot or cold high pressure homogenization technique [4]. They are more stable than liposomes because of their rigid core consisting of hydro- phobic lipids which are solid at room and body temperature, surrounded by a monolayer of phospholipids. They can be stabilized by high level of sur- factants. They are less toxic than polymeric nanoparticles because of ease of biodegradation. Solid lipid nanoparticles have important advantages, such as their physiological compounds and large scale production favoured their feasibility thus avoiding organic solvents in the manufacturing process [11]. Nanosuspensions are poor water soluble drugs dispersed in aqueous phase containing stabilizing agent [4]. Clofazimine, a riminophenazine compound used for treating patients with M. avium infection and because of its poor solubility the drug usage was restricted, now to overcome this problem of solubility. It was formulated as a nanosuspension (385 nm) and was administered to mice intravenously which has resulted in reduction of bacterial loads in the liver, lungs, spleen of mice [12].