Capacitive Biosensors and Molecularly Imprinted Electrodes

Gizem Ertürk; Bo Mattiasson

doi:10.3390/s17020390

Capacitive Biosensors and Molecularly Imprinted Electrodes

Ertürk, Gizem; Mattiasson, Bo 2017-02-17 00:00:00 sensors Review Capacitive Biosensors and Molecularly Imprinted Electrodes 1 , 1 , 2 Gizem Ertürk * and Bo Mattiasson CapSenze Biosystems AB, Lund 223 63, Sweden; [email protected] Department of Biotechnology, Lund University, Lund 222 40, Sweden * Correspondence: [email protected] Academic Editor: Nicole Jaffrezic-Renault Received: 14 December 2016; Accepted: 8 February 2017; Published: 17 February 2017 Abstract: Capacitive biosensors belong to the group of afﬁnity biosensors that operate by registering direct binding between the sensor surface and the target molecule. This type of biosensors measures the changes in dielectric properties and/or thickness of the dielectric layer at the electrolyte/electrode interface. Capacitive biosensors have so far been successfully used for detection of proteins, nucleotides, heavy metals, saccharides, small organic molecules and microbial cells. In recent years, the microcontact imprinting method has been used to create very sensitive and selective biorecognition cavities on surfaces of capacitive electrodes. This chapter summarizes the principle and different applications of capacitive biosensors with an emphasis on microcontact imprinting method with its recent capacitive biosensor applications. Keywords: capacitive biosensors; afﬁnity biosensors; microcontact imprinting 1. Introduction Afﬁnity biosensors can be divided into two main groups: those that measure direct binding between the target molecule and the afﬁnity surface on the sensor, and those biosensors which are adopted to binding assays using labelled reagents [1]. Biosensors operating with labelled afﬁnity-reagents are variations of conventional immunoassay technology in which ﬂuorescent markers, active enzymes, magnetic beads, radioactive species or quantum dots are generally used as labelling agents to label the target molecules [2,3]. Labelling is generally used to signiﬁcantly facilitate the signal generation and to conﬁrm the interaction between the probe and target molecules [4]. An important feature in using labelled reagents is the ampliﬁcation of the registered signals and thereby also the sensitivity one can reach. Assays based on the use of labelled reagents are time consuming and the labelled reagents are often expensive. Furthermore, assays with labelled reagents are usually multistep processes and that limit their application for real-time measurements. The design of label-free afﬁnity based biosensors is the objective of much current research. The aims are to establish alternative methods to the commercial ELISA-based immunoassays. The most attractive features of these types of biosensors are that they allow the monitoring of the analytes directly and in real-time [5]. Biosensors can be divided into four main groups according to transducer types [6]. These main groups involve: electrochemical transducers which involve potentiometric, voltammetric, conductometric, impedimetric and ﬁeld-effect transistors; optical transducers which include surface plasmon resonance (SPR) biosensors; piezoelectric transducers to which quartz crystal microbalance (QCM) biosensors can be given as an example; and thermometric transducers which measure the amount of heat with a sensitive thermistor to determine the analyte concentration. Sensors 2017, 17, 390; doi:10.3390/s17020390 www.mdpi.com/journal/sensors Sensors 2017, 17, 390 2 of 21 Among the different types of label-free biosensors, electrochemical biosensors have received particular attention owing to their properties [4]. These biosensors can also be miniaturized which is very important for many applications that need portable integrated systems. Miniaturization not only allows use at point of care, in a clinic, doctor ’s ofﬁce or at home but also reduces the cost of the diagnostic assays. Electrical biosensors fulﬁl these purposes as being fast, cheap, portable, miniaturized and label-free devices. Electrical biosensors can be classiﬁed as amperometric, voltametric impedance or capacitive sensors. This review deals with capacitive biosensors, different applications of capacitive biosensors developed for both detection of various targets and by using molecular imprinting technology with an emphasis on microcontact imprinting method. In this aspect, this is a novel review which includes the two technologies, capacitive biosensors and molecular imprinting; at the same time with lots of examples from published reports. The reports in molecular imprinting section were selected mainly from capacitive biosensors developed by using microcontact imprinting method. 2. Capacitive Biosensors Capacitive biosensors belong to the sub-category of impedance biosensors [3]. Capacitive biosensors measure the change in dielectric properties and/or thickness of the dielectric layer at the electrolyte-electrode interface when an analyte interacts with the receptor which is immobilized on the insulating dielectric layer [1]. The electric capacitance between the working electrode (an electrolytic capacitor/the ﬁrst plate) and the electrolyte (the second plate) is given by Equation (1) [7]: C = (" "A)/d (1) where " is the dielectric constant of the medium between plates, " is the permittivity of the free space 12 2 (8.85 10 F/m), A is the surface area of the plates (m ) and d is the thickness of the insulating layer (m). According to the given equation above, when the distance between the plates increases, the total capacitance decreases. In other words, in the assaying principle of this type of capacitive biosensors when a target molecule binds to the receptor, displacement of the counter ions around the capacitive electrode results in a decrease in the capacitance. The higher the amount of target molecules bound to the receptor is, the greater is the achieved displacement and the decrease in the registered capacitance [8]. The assaying principle of capacitive biosensors which are developed according to this rule is shown in Figure 1. The Equation (1) can be represented by two capacitors in series where the inner part includes the dielectric layer (C ) and the outer one corresponds to the biomolecule layer (C ). Then, the total dl bm capacitance (C ) can be described as Equation (2) [7]. 1 1 1 = + (2) C C C dl bm The electrochemical capacitors which are described based on the above-mentioned equation are known as constant phase element (CPE). The presence of CPE indicates that the observed capacitance of the system is frequency dependent. Capacitance can also be deﬁned as Equation (3) [7]. Z = (3) !C where Z is the impedance and ! is the radial frequency expressed in rads . This model implies that all of the measured current is capacitive. Sensors 2017, 17, 390 3 of 21 Sensors 2017, 17, 390 3 of 21 Figure 1. (A) Schematic diagram showing the change in capacitance (ΔC) as a function of time when Figure 1. (A) Schematic diagram showing the change in capacitance (DC) as a function of time when the analyte (IgG) interacts with the receptor molecule (Protein A) immobilized on the surface of the the analyte (IgG) interacts with the receptor molecule (Protein A) immobilized on the surface of the electrode. Subsequent rise in signal is due to the dissociation after the injection of the regeneration electrode. Subsequent rise in signal is due to the dissociation after the injection of the regeneration solution. In an ideal sensorgram, the baseline should turn back to the original level after regeneration solution. In an ideal sensorgram, the baseline should turn back to the original level after regeneration of of the surface; (B) Immobilization of the receptor molecule on the transducer surface via a the surface; (B) Immobilization of the receptor molecule on the transducer surface via a self-assembled self-assembled monolayer (SAM) of alkylthiols. When the target molecule interacts with the receptor, monolayer (SAM) of alkylthiols. When the target molecule interacts with the receptor, this creates this creates a double layer of counter ions around the gold transducer which results in a change in the a double layer of counter ions around the gold transducer which results in a change in the capacitance. capacitance. (Reproduced from Reference [8] with permission). (Reproduced from Reference [8] with permission). 3. Different Applications of Capacitive Biosensors 3. Different Applications of Capacitive Biosensors Different applications of capacitive biosensors developed for different targets are summarized Different applications of capacitive biosensors developed for different targets are summarized in in Table 1. Table 1. Sensors 2017, 17, 390 4 of 21 Table 1. Different applications of capacitive biosensors developed for different targets. Limit of Target Sensor Preparation Method Dynamic range (M) Selectivity Stability Ref. Detection (M) Immobilization of anti-CT antibodies on self-assembled monolayer (SAM) of 13 10 14 Cholera toxin (CT) 1.0 10 –1.0 10 1.0 10 uiu N/D N/D [9] lipoic acid and 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) Immobilization of anti-CT on Up to 36 times 18 12 20 Cholera toxin (CT) gold nanoparticles incorporated 0.1 10 –10 10 9.0 10 N/D [10] with an RSD of 2.5% on a poly-tyramine layer Proteins Immobilization of anti-HIV 1 p24 antigen on gold nanoparticles 20 17 20 HIV-1 p24 antigen 10.1 10 –10.1 10 3.32 10 N/D N/D [11] incorporated on a poly-tyramine layer Immobilization of anti-VEGF aptamer ﬁrst capturing the VEGF 14 11 VEGF protein then, sandwiching with N/D N/D N/D [12] 13 10 –2.6 10 antibody-conjugated magnetic beads Oligo-T was used as the competing agent, when the Covalent attachment of 25-mer temperature was increased 8 11 11 25-mer oligo C oligo C on poly-tyramine 10 –10 10 N/D [13] from RT to 50 C, the DC modiﬁed electrode value decreased from 2 2 48 nFcm to 3 nFcm GOPTS functionalized Thiol modiﬁed oligonucleotides surfaces were more stable at were immobilized on Au and 4 C. Ten-fold decrease in 6 3 ssDNA N/D N/D [14] 0.5 10 –1.0 10 Nucleic acids 3-glycidoxypropyl-tri-methoxy ﬂuorescence intensity after silane (GOPTS) 1 week even when the substrates were stored at 4 C. Immobilization of ssDNA 10 11 Nampt aptamers on SAM of 0–45 10 1.8 10 N/D N/D [15] mercaptopropionic acid (MPA) Complementary DNA Immobilization of pyrrolidinyl Could be reused for provided a much higher DC 11 10 12 Target DNA peptide nucleic acid probes 1.0 10 –1.0 10 6–10 10 58–73 times with an average [16] compared to single and (acpcPNA) residual activity of 98% double mismatched DNA Sensors 2017, 17, 390 5 of 21 Table 1. Cont. Limit of Target Sensor Preparation Method Dynamic range (M) Selectivity Stability Ref. Detection (M) Based on the interaction For the ﬁrst 35 cycles, the between E. coli and concanavalin residual activity was 1 6 1 1 Total bacteria 12 CFUmL –1.2 10 CFUmL 12 CFUmL N/D [17] A immobilized on a modiﬁed 95% 3% (RSD = 3.2%). Cells gold surface After 35 cycles, it was 85%. E. coli cells immobilized on SAM 5 1 7 1 E. coli N/D N/D N/D [18] 8 10 CFUmL –8 10 CFUmL of Mercaptopropionic acid (MPA) Immobilization of metal Hg(II), Cu(II), 15 3 resistance and metal regulatory 10 –10 N/D N/D N/D [19] Zn(II), Cd(II) proteins on gold electrode 1. Immobilization of whole bacterial cell to emit Heavy metals 84% of the activity loss a bioluminescent/ﬂuorescent 0–200 10 within 6 days N/D [20] 1.0 10 Cu(II), Cd(II), signal in the presence of heavy Hg(II) metal ions 2. Immobilization of heavy metal 15 1 10 –10 Stable over 16 days binding proteins Small sugars including D-fructose, D-mannose, Immobilization of ConA on gold A neglectable loss in D-maltose, 6 2 6 Glucose nanoparticles incorporated on the 1.0 10 –1.0 10 1.0 10 sensitivity after [21] Saccharides methyl--D-glucopyranoside, tyramine modiﬁed gold electrode 10 cycles (7.5%) methyl--D-mannopyranoside also bound instead of glucose Immobilization of ConA and Small molecules and high 5 1 6 Glucose replacement of small glucose 1.0 10 –1.0 10 1.0 10 molecular weight dextran N/D [22] with the large glucose polymer also bound instead of glucose Immobilization of molecularly Cross reactant contribution 6 6 Metergoline imprinted spherical beads on 1–50 10 1.0 10 N/D [23] was maximum 1.3 nF modiﬁed gold electrode Competing agents’ binding Little variation over 28 6 9 12 Aﬂatoxin B1 Bioimprinting 3.2 10 –3.2 10 6.0 10 was signiﬁcantly lower than injections with non-reduced [24] Small molecules aﬂatoxin B1 Schiff’s bases Monoclonal anti-OTA Differences for ochratoxin B Ochratoxin A immobilization on Si N 3 4 12 12 2.47–49.52 10 4.57 10 and aﬂatoxin B1 were N/D [25] (OTA) substrate combined with not signiﬁcant magnetic nanoparticles (MNPs) Sensors 2017, 17, 390 6 of 21 3.1. Protein Detection Labib et al. [9] developed a sensitive method for detection of cholera toxin (CT) using a ﬂow-injection capacitive immunosensor based on self-assembled monolayers. Monoclonal antibodies against the subunit of CT (anti-CT) were immobilized on the gold electrode surface which was modiﬁed by lipoic acid and 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC). The immunosensor showed linear response to CT concentrations in the concentration range between 1.0 10 M and 10 14 1.0 10 M under optimized conditions. Limit of detection (LOD) value was 1.0 10 M. The LOD value obtained from capacitive immunosensor was compared with the LOD values that were obtained from sandwich ELISA and SPR based immunosensors. The ELISA had a LOD of 12 11 1.2 10 M whereas SPR had a LOD value of 1.0 10 M. The results proved that the method is more sensitive than the other two techniques used in this study. In another study by the same research group, a label-free capacitive immunosensor was developed for direct detection of CT present at sub-attomolar level. Gold nanoparticles (AuNPs) were incorporated on a polytyramine modiﬁed gold electrode and anti-CT antibody was immobilized on this surface. Tyramine provides free amino groups which are very useful to immobilize the afﬁnity ligand to the transducer. At the same time very thin and uniform ﬁlms can be formed on the electrode surfaces during the electro-polymerization of tyramine. After the immobilization step, the formation of antigen-antibody complexes resulted in a change in capacitance and by this way the concentration of CT was determined. The dynamic range was between 0.1 aM and 10 pM where the LOD value was 9 10 M (0.09 aM). The electrode could be regenerated with a good reproducibility for up to 36 times with a relative standard deviation (RSD) value of 2.5%. Real sample analyses were performed from water samples collected from a local stream and matrix effect was eliminated with a 10.000 times dilution prior to analysis. The developed system had potential to be used as a portable electrochemical analyser for ﬁeld conditions [10]. The same strategy was used to develop a capacitive biosensor for sensitive detection of HIV-1 p24 antigen [11]. Following polytyramine electro-polymerization on the gold electrode surface, gold nanoparticles were incorporated onto the electrode and then, anti-HIV-1 p24 monoclonal antibodies were immobilized on top. HIV-1 p24 antigen was detected from standard p24 solutions in the 20 17 concentration range of 10.1 10 to 10.1 10 M. The reasons for extreme sensitivity of the capacitive biosensors were explained by the authors in two ways. Firstly, the capacitive technique is very convenient to detect the size and position of the electrical double layer formed at the interface. However, in other electrochemical techniques such as the amperometric technique, the detection is based on the measurement of only the transport of electrons. The second factor is the increased surface area on the electrode via the use of immobilized AuNPs. The gold nanoparticles also signiﬁcantly contributed to the decrease in the interfacial resistance which facilitated the electron transfer at the electrode surface [11]. Qureshi et al. [12] developed a capacitive aptamer based sensor for detection of vascular endothelial growth factor (VEGF) in human serum. Systematic evolution of ligands by exponential enrichment (SELEX) process was utilized to select the highly speciﬁc and selective anti-VEGF aptamer to bind VEGF to the aptasensor surface. When a sandwich assay was tested by forming sandwich complex with anti-VEGF aptamer+VEGF+anti-VEGF antibody, the generated signal was enhanced by 3–8 folds compared to the direct assay. The developed sensor showed a dynamic detection range 14 11 from 13 10 M to 2.6 10 M of VEGF protein in human serum. The results showed that the developed system could be successfully used in clinical diagnosis to detect biomarkers in real samples in a convenient and sensitive way. 3.2. Nucleic Acid Detection Mahadhy et al. [26] used in a model study the capacitive sensor to monitor the capture of complementary single-stranded nucleic acids. A 25-mer oligo-C was immobilized onto the polytyramine modiﬁed gold electrode surface. Temperature was raised to 50 C to reduce non-speciﬁc Sensors 2017, 17, 390 7 of 21 hybridization in order to increase the selectivity and hybridization was used in order to amplify the signal by using longer nucleic acid molecules. Later, Mahadhy et al. [13] developed a promising ultrasensitive, automated ﬂow-based and portable gene sensor. The PCR-free biosensor proved the possibility for powerful detection of foodborne pathogens in diagnostic situations and multi-drug resistant bacteria in the near future. Rapid detection of foodborne pathogens is crucial before many people are infected while early detection of multi-drug resistant bacteria is important to isolate the infected patients earlier and to reduce the risk of spreading. Two functionalization layers; gold (Au) and 3-glycidoxypropyl-tri-methoxy silane (GOPTS) were used to immobilize thiol modiﬁed oligonucleotides on silicon surfaces. GOPTS showed better performance as a functionalization layer because the hybridization efﬁciency was higher, the stability over time was better and regeneration of the surface after analyte binding was easier. Therefore for the development of microcantilever or micro-membrane based biosensors, GOPTS might be a more promising alternative [14]. A single stranded DNA (ssDNA) aptamer was developed to bind to nicotinamide phosphoribosyl transferase (Nampt) through SELEX and implemented in a capacitive biosensor [15]. The LOD for 11 10 Nampt was 1.8 10 M with a dynamic detection range in serum of up to 9 10 M. Nampt is an important biomarker for obesity-related metabolic diseases, some types of cancers and chronic 1 11 diseases. Normal level of Nampt in human plasma is around 15 ngmL (27 10 M). Therefore, the developed system has potential as a diagnostic tool and in point-of-care applications. Pyrrolidinyl peptide nucleic acid probes were immobilized onto the self-assembled monolayer (SAM) modiﬁed gold electrode surface to develop a DNA capacitive biosensor [16]. Four different alkanethiols with various chain lengths were used as a SAM to determine the inﬂuence of the length and the terminating head group of blocking thiols on the sensitivity and speciﬁcity. In the study, the blocking thiol which had an equal length to the –OH terminating head group gave the highest sensitivity and high binding speciﬁcity. Thus far, there are no good examples on use of MIPs in connection to monitoring of speciﬁc DNA sequences. Since the potential is large, one can expect such developments to happen in the near future. 3.3. Cell Detection Jantra et al. [17] developed a label-free afﬁnity biosensor for detecting and enumerating total bacteria based on the interaction between E. coli and Concanavalin A (Con A) immobilized on a modiﬁed gold surface. The analyses were completed in less than 20 min with both SPR and 7 1 impedimetric capacitive biosensor. Compared to SPR (LOD: 6.1 10 CFUmL ), the capacitive system showed much higher sensitivity (LOD: 12 CFUmL ). The developed system might be used successfully for total bacteria analysis from water sources. Rydosz et al. [27] developed a new type of label-free microwave sensor in a form of interdigitated capacitor for bacterial lipopolysaccharide detection. The sensor surface was coated with T4 phage gp37 adhesin. The adhesin molecule bound E. coli by recognizing its bacterial host lipopolysaccharide (LPS). The binding was highly speciﬁc and irreversible. Recognition between the phage adhesion and bacterial LPS was based on the recognition of saccharide determinants of LPS which means very speciﬁc determination of bacterial strain or its endotoxins within the genus and the species. The selectivity experiments showed that the response for speciﬁc LPSs was signiﬁcantly different from the reference measurements and the response for the non-speciﬁc LPSs was very close to the reference values. The developed method was promising for label-free LPS detection and could be used as an alternative for ﬁber-optic, electrochemical and classic biochemical and immunochemical methods. In their other study [28], same authors used bacteriophage-adhesin-coated long-period gratings for recognition of bacterial lipopolysaccharides. Long-period gratings (LPG) bio-functionalization methodology was based on coating the LPG surface with nickel ions which were capable of binding of gp37-histidine tag. The advantage of using adhesins for the bio-functionalization of the biosensor was to give ability Sensors 2017, 17, 390 8 of 21 for low-cost bio-sensitive molecule exchange and surface regeneration. In this work, for the ﬁrst time, adhesion has been applied for bacteria and their endotoxin detection. T4 phage adhesin bound E. coli B LPS in its native or denatured form in a highly speciﬁc and irreversible way. Rocha et al. [29] used alternating current electrokinetics (ACEK) capacitive sensing to detect and quantify the microbial cell abundance in aquatic systems. Microbial abundance was detected by measuring the electrical signal. Three different microbial cell cultures including Bacillus subtilis, Alcanivorax borkumensis and Microcystis aeruginosa were detected by using the developed system. The results showed that the sensor is capable of reliably detecting microbial cells even though they have major physiological differences between Gram-positive (B. subtilis), Gram-negative (A. borkumensis) and cyanobacteria (M. aeruginosa). The system is promising to detect and estimate microorganism population sizes in batch cultures, environmentally sourced seawater and groundwater systems. For expanded use and commercialization of nanotechnology products, toxicity determination is an important ﬁeld application. For this purpose, Qureshi et al. [18] developed a whole-cell based capacitive biosensor to determine the biological toxicity of nanoparticles (NPs). They used iron oxide (Fe O ) nanoparticles as models in the study. The living E. coli cells were immobilized on the capacitive 3 4 sensor chips. Then, these chips were interacted with different sizes of Fe O NPs (5, 20 and 100 nm). 3 4 The smallest Fe O NPs resulted in a maximum capacitance change because they were able to interact 3 4 with E. coli cells on the sensor chip very efﬁciently. The morphological changes on the surface of E. coli cells after interacting with Fe O NPs were examined with SEM. 3 4 3.4. Heavy Metal Detection Two metal binding proteins were over-expressed in E. coli, puriﬁed and immobilized on a thiol-modiﬁed capacitive sensor surface. Capacitive sensor was used to monitor conformational changes following heavy metal binding including copper, cadmium, mercury and zinc. Metal ion detection could be done in down to femtomolar concentrations with the developed system [30]. The same group expressed and puriﬁed metal resistance and metal regulatory proteins from bacterial strains and immobilized these proteins on the capacitive biosensor surface for heavy metal detection [19]. The system allowed the detection of heavy metals including Hg(II), Cu(II), Zn(II) and Cd(II) in pure solutions down to 10 M concentrations. Corbisier at al. [20] established two different biosensor technologies for detection of several heavy metal ions in environmental samples. The principle of the ﬁrst approach was to develop whole cell bacterial biosensors which emitted a bioluminescent/ﬂuorescent signal in the presence of heavy metal ions. In the second approach, direct interaction between metal binding proteins and heavy metal ions was used as the detection principle in the capacitive biosensors. In the study, the main advantage of the whole cell based sensors was their ability to react only to biologically available metal ions, whereas the latter one (protein based sensors) was more sensitive towards metal ions. MIPs selective for heavy metal ions have been presented in connection to separation technology [31–36]. It is an obvious development that also sensors for heavy metal ions will be developed based on selective MIPs. 3.5. Saccharide Detection By using capacitive biosensors, Labib et al. [21] used the capacitive biosensor to detect glucose based on gold nanoparticles which were ﬁxed on a poly-tyramine modiﬁed gold electrode surface. Dextran (MW: 39 kDa) was used as a regeneration agent by utilizing a competitive assay for glucose in 6 2 the study. The dynamic range for glucose detection was between 1.0 10 and 1.0 10 M with a LOD value of 1.0 10 M. By using capacitive biosensors, the same authors [22] developed a technique based on the competition between a small molecular mass analyte and a large analyte-carrier conjugate. In the basis of the technique, when a large glucose polymer binds to the biorecognition molecule (Con A) immobilized on the electrode surface, it results in a decrease in the capacitance. Then, in the Sensors 2017, 17, 390 9 of 21 next step, when the low molecular mass analyte (glucose) is introduced to the system, the effect is reverse, the small glucose molecule will replace the large glucose polymer which is bound on the immobilized Con A as shown in Figure 2. By measuring the shift-back in capacitance, the glucose concentration could be determined by the technique. The authors used this technique to measure IgG as a glycoconjugate and detect its aggregation using immobilized Con A. When the glycoconjugate (IgG) was injected, the decrease in capacitance was measured to determine its concentration. In the second step, when concentrated glucose was injected into the system, the increase in capacitance was employed to determine the glucose concentration. The results showed that this technique is promising for monitoring Sensors 2017, 17 small , 390 molecules with high sensitivity and broad detection range. 10 of 21 Figure 2. Schematic representation of the competitive glucose binding assay. (a) When glucose is Figure 2. Schematic representation of the competitive glucose binding assay. (a) When glucose is injected into the capacitive system, it binds to the immobilized Concanavalin A (ConA) on the surface. injected into the capacitive system, it binds to the immobilized Concanavalin A (ConA) on the surface. However, this binding does not make any change in the capacitance level, as shown in the graph on However, this binding does not make any change in the capacitance level, as shown in the graph on the the right, due to the small size of the glucose molecule; (b) When a glucose polymer (dextran) is right, due to the small size of the glucose molecule; (b) When a glucose polymer (dextran) is injected injected into the system, binding of this big polymer to ConA results in a decrease in the capacitance into the system, binding of this big polymer to ConA results in a decrease in the capacitance signal; signal; (c) When glucose is injected into the system again, displacement of dextran with glucose results (c) When glucose is injected into the system again, displacement of dextran with glucose results in the in the capacitance turn back to the original baseline level. (Reproduced from Reference [22] with capacitance turn back to the original baseline level. (Reproduced from Reference [22] with permission). permission). 3.6. Small Organic Molecules The radio frequency (RF) detection method is one of the promising methods for glucose detection out of several detection methods available [37]. When an analyte is injected into the RF biosensor, Small molecules such as pesticides, herbicides, and antibiotics are widely discarded and encountered in naturally flowing waters. These pollutants in environment have an impact on changes will occur owing to the inductive and capacitive effects [38]. These changes will cause communities and eco-systems. Therefore, detection of these molecules in a sensitive, cheap, robust losses and considerable shifts in the resonance frequency of the device. The change in capacitance is and fast way is crucial. Lenain et al. [23] chose metergoline as a model compound representingsmall proportional to the dielectric constant and the distance between the biomolecule layer and the dielectric organic molecules such as pharmaceutical residues. Emulsion polymerization was used to produce layer. A reusable robust RF biosensor was developed by Kim et al. [38] to monitor real-time glucose small, uniformly sized, spherical MIPs. These MIP beads were attached to the poly-tyramine level in human serum. The resonance behaviour of the system was analysed with human serum modified gold surface. Scanning electron microscope (SEM) images of the electrode surface are 1 1 samples containing different glucose concentrations ranging from 148–268 mgdL , 105–225 mgdL −6 −6 shown in Figure 3. Working range for metergoline was from 1.0 × 10 M to 50 × 10 M with a LOD and at a deionized water glucose concentration in the range of 25–500 mgdL . The response time for −6 value of 1.0 × 10 M. In cross reactivity analysis, even though the structural analogs showed binding, glucose was measured as 60 s with a LOD value of 8.01 mgdL . A total of 21 different experiments this contribution was only around 1.3 nF. The sensor response was more stable at higher ionic for each strength but the extent of concentration of serum ca and paci D-gl tance cha ucosensolution ge for different co were analysed ncentrations for reusability of analyte w and the as less relative pronounced compared to lower electrolyte concentrations. standard deviation (RSD) was less than 1% for each concentrations of serum samples and aqueous Bioimprinting is a technology used to mimic specific sites for modification of biological D-glucose solutions. molecules. The process consists of four steps, as shown in Figure 4 [24]. In the area of MIPs used for bioseparation, much work has been done concerning carbohydrates (1) Unfold [39 ing – t 41 he conform ]. It is obvious ation of t that he st one arting prot can make ein un good der acid MIPs ic con with ditions high ; efﬁciency in binding (2) Addition of template molecule and allow interaction between the template molecule and the denatured protein in order to form new molecular configurations; (3) Cross-linking of the protein to stabilize the new molecular protein conformation; and (4) Dialysis to remove the template molecule. Sensors 2017, 17, 390 10 of 21 target molecules or fragments thereof. Based on the observations from afﬁnity chromatography, one can foresee that such systems will also soon be presented for capacitive biosensors [42–48] (Table 2). Table 2. Molecularly imprinted polymers (MIPs) produced with high binding efﬁciency for afﬁnity chromatography applications. Template Method Matrix Comments Ref. Preconcentration on BAP BAP-imprinted poly with HPLC equipped with Benzo[a]pyrene (2-hydroxyethylmethacrylate- Aqueous solutions [42] a ﬂuorescence detector (BAP) N-methacryloyl-(L)-phenylalanine (HPLC-FLD) composite cryogel cartridge MIP-solid phase extraction Melamine imprinted Extraction and enrichment Melamine Water + milk [43] monolithic cartridges of melamine Cholesterol imprinted Gastrointestinal Cholesterol adsorption Cholesterol [44] polymeric nanospheres mimicking solution Iron chelated poly Catalase puriﬁcation (2-hydroxyethylmethacrylate- Catalase Rat liver [45] from rat liver N-methacryloyl-(L)-glutamic acid cryogel discs Chiral separation of l-phenylalanine with L-phenylalanine L-Phe imprinted Aqueous solutions [46] FPLC (fast protein (L-Phe) cryogel cartridges liquid chromatography) Separation of triazine with Triazine imprinted capillary electro- Triazine Aqueous solutions [47] monolithic columns chromatography (CEC) Cytochrome c puriﬁcation Surface imprinted bacterial Cytochrome c Rat liver [48] from rat liver cellulose nanoﬁbers 3.6. Small Organic Molecules Small molecules such as pesticides, herbicides, and antibiotics are widely discarded and encountered in naturally ﬂowing waters. These pollutants in environment have an impact on communities and eco-systems. Therefore, detection of these molecules in a sensitive, cheap, robust and fast way is crucial. Lenain et al. [23] chose metergoline as a model compound representingsmall organic molecules such as pharmaceutical residues. Emulsion polymerization was used to produce small, uniformly sized, spherical MIPs. These MIP beads were attached to the poly-tyramine modiﬁed gold surface. Scanning electron microscope (SEM) images of the electrode surface are shown in Figure 3. 6 6 6 Working range for metergoline was from 1.0 10 M to 50 10 M with a LOD value of 1.0 10 M. In cross reactivity analysis, even though the structural analogs showed binding, this contribution was only around 1.3 nF. The sensor response was more stable at higher ionic strength but the extent of capacitance change for different concentrations of analyte was less pronounced compared to lower electrolyte concentrations. Bioimprinting is a technology used to mimic speciﬁc sites for modiﬁcation of biological molecules. The process consists of four steps, as shown in Figure 4 [24]. (1) Unfolding the conformation of the starting protein under acidic conditions; (2) Addition of template molecule and allow interaction between the template molecule and the denatured protein in order to form new molecular conﬁgurations; Sensors 2017, 17, 390 11 of 21 (3) Cross-linking of the protein to stabilize the new molecular protein conformation; and (4) Dialysis to remove the template molecule. Sensors 2017, 17, 390 11 of 21 Sensors 2017, 17, 390 11 of 21 c d c d Figure 3. Scanning electron microscope (SEM) pictures of the electrode surface after functionalization Figure 3. Scanning electron microscope (SEM) pictures of the electrode surface after functionalization Figure 3. Scanning electron microscope (SEM) pictures of the electrode surface after functionalization with imprinted polymers. From left to right, top to bottom: (a) SEM picture of electrode surface; (b,c) with imprinted polymers. From left to right, top to bottom: (a) SEM picture of electrode surface; with imprinted polymers. From left to right, top to bottom: (a) SEM picture of electrode surface; (b,c) SEM pictures of centre of the electrode; and (d) SEM picture of the border between the gold layer and (b,c) SEM pictures of centre of the electrode; and (d) SEM picture of the border between the gold layer SEM pictures of centre of the electrode; and (d) SEM picture of the border between the gold layer and wafer. (Reproduced from Reference [23] with permission). and wafer. (Reproduced from Reference [23] with permission). wafer. (Reproduced from Reference [23] with permission). Figure 4. Schematic representation of bio-imprinting process. (Reproduced from Reference [24] with Figure 4. Schematic representation of bio-imprinting process. (Reproduced from Reference [24] with permission). permission). Figure 4. Schematic representation of bio-imprinting process. (Reproduced from Reference [24] with permission). Gutierrez et al. [24] used bioimprinting to develop a capacitive biosensor for aflatoxin detection. Gutierrez et al. [24] used bioimprinting to develop a capacitive biosensor for aﬂatoxin detection. Aflatoxins are natural food contaminants with a high risk for human health. Ovalbumin was used as Gutierrez et al. [24] used bioimprinting to develop a capacitive biosensor for aflatoxin detection. Aﬂatoxins platform ar fo e natural r bioimprintin food contaminants g of aflatoxin because with a when bovine serum albumi high risk for human health. n (BSA) w Ovalbumin as used, there was used Aflatoxins are natural food contaminants with a high risk for human health. Ovalbumin was used as was no change in capacitance owing to the high hydrophobicity of sites of BSA. Three competitive as platform for bioimprinting of aﬂatoxin because when bovine serum albumin (BSA) was used, platform mycotoxi for bioimprintin ns were used in the cross-rea g of aflatoxin because ctivity a when bovine serum albumi nalysis and the changes inn ca (BSA) w pacitaance were s used, there there was no change in capacitance owing to the high hydrophobicity of sites of BSA. Three competitive significantly lower than that registered from aflatoxin solution. was no change in capacitance owing to the high hydrophobicity of sites of BSA. Three competitive mycotoxins were used in the cross-reactivity analysis and the changes in capacitance were signiﬁcantly mycotoxins were used in the cross-reactivity analysis and the changes in capacitance were lower than that registered from aﬂatoxin solution. significantly lower than that registered from aflatoxin solution. Sensors 2017, 17, 390 12 of 21 Sensors 2017, 17, 390 12 of 21 Silicon nitride substrate (Si N ) combined with magnetic nanoparticles (MNPs) was used to 3 4 develop a capacitive immunosensor for ochratoxin A (OTA) detection. Silicon nitride allows an easy Silicon nitride substrate (Si3N4) combined with magnetic nanoparticles (MNPs) was used to control of the ﬁlm composition and thickness and also prevents the undesirable impurities. These are develop a capacitive immunosensor for ochratoxin A (OTA) detection. Silicon nitride allows an easy the main advantages of the substrate used in the study. Magnetic nanoparticles comprised of control of the film composition and thickness and also prevents the undesirable impurities. These are a conductive core and a carboxylic acid modiﬁed shell which was used to immobilize OTA antibodies. the main advantages of the substrate used in the study. Magnetic nanoparticles comprised of a The LOD value was calculated as 4.57 10 M in the study and the selectivity results against conductive core and a carboxylic acid modified shell which was used to immobilize OTA antibodies. ochratoxin B and aﬂatoxin G1 showed that −the 12 potential difference was not so signiﬁcant when The LOD value was calculated as 4.57 × 10 M in the study and the selectivity results against compared to the difference for OTA detection [25]. ochratoxin B and aflatoxin G1 showed that the potential difference was not so significant when compared to the difference for OTA detection [25]. 4. Molecular Imprinting 4. Molecular Imprinting During the early 1970s, Wulff et al. [49] and Klotz et al. [50] introduced the molecular imprinting to imprint templates in organic polymers. Then, Mosbach et al. [51] reported the use of molecularly During the early 1970s, Wulff et al. [49] and Klotz et al. [50] introduced the molecular imprinting imprinted polymers (MIPs) in biosensors instead of antibodies which was a breakthrough. to imprint templates in organic polymers. Then, Mosbach et al. [51] reported the use of molecularly imprint The formation ed polymers of ( MIPs MIPs) involves in biosensor three s in steps: stead of antibodies which was a breakthrough. The formation of MIPs involves three steps: (1) Pre-complexation of functional monomers around the template molecule in solution either by (1) Pre-complexation of functional monomers around the template molecule in solution either by forming covalent bonds or by self-assembling with non-covalent bonds; forming covalent bonds or by self-assembling with non-covalent bonds; (2) Polymerization of the resulting complex in the presence of cross-linking monomers and suitable (2) Polymerization of the resulting complex in the presence of cross-linking monomers and suitable solvents/ionic liquids as porogens; and solvents/ionic liquids as porogens; and (3) Removal of template molecule from the synthesized polymer. (3) Removal of template molecule from the synthesized polymer. The The resu resulting lting MIP MIP cont contains ainsr recogn ecognition ition cavities cavities c capable apable of ofsel selective ective recogn recognition ition of compound of compounds s that fit these cavities with respect to shape, size, position and orientation of the recognition sites [52]. that ﬁt these cavities with respect to shape, size, position and orientation of the recognition sites [52]. How MIPs can mimic natural recognition units in different applications are shown schematically in How MIPs can mimic natural recognition units in different applications are shown schematically in Figure 5. Figure 5. Figure 5. Different applications of MIPs in: (A) immunosensors; (B) enzyme-linked immunosorbent Figure 5. Different applications of MIPs in: (A) immunosensors; (B) enzyme-linked immunosorbent assay (ELISA); (C) enzyme electrodes, reaction rate and analyte concentration of enzyme electrodes assay (ELISA); (C) enzyme electrodes, reaction rate and analyte concentration of enzyme and catalytic MIP-coated electrodes can be estimated by electroactive substrate/product electrodes and catalytic MIP-coated electrodes can be estimated by electroactive substrate/product consumption/production during the catalytic reaction or electron transfer from the electrode surface consumption/production during the catalytic reaction or electron transfer from the electrode surface to to the active centre of enzyme/MIP; (D) DNA chips; and (E) enzyme immobilization and competitive the active centre of enzyme/MIP; (D) DNA chips; and (E) enzyme immobilization and competitive binding of the analyte. (Reproduced from Reference [52] with permission). binding of the analyte. (Reproduced from Reference [52] with permission). MIP technology has successfully been used for imprinting of low molecular weight templates. However there are still some difﬁculties of molecular imprinting technique when it is used for Sensors 2017, 17, 390 13 of 21 macromolecular templates including proteins. Due to this, many researchers have focused on the alternative techniques including imprinting the template directly onto a substrate or immobilizing the template protein on a glass support and use it as a protein stamp. The latter is called microcontact imprinting. 5. Microcontact Imprinting Microcontact imprinting technique was ﬁrst introduced by Chou et al. [53]. In the study, the authors formed the microcontact imprints between two cleaned glass surfaces. Template protein was immobilized on the cover slip and then, functional monomer was added on top in order to allow site-speciﬁc organization of the functional monomer by the template. In the next step, a drop of solution including cross-linker and initiator was dropped on the pre-modiﬁed support glass. Support glass was brought into contact with the cover slip to provide the two functionalized surfaces to contact. The glass assemblies were placed in a UV reactor to initiate the polymerization and it continued for 17 h. After polymerization, the cover slip was removed from the surface with forceps and the support was washed with various solutions. Polymer ﬁlm thickness measurements showed it to be around 10 m. The imprinted polymers showed better selectivity for their native templates than competing proteins. By this study, it was reported that microcontact imprinting technique might be used successfully for relatively large proteins in biosensor applications for detection and quantiﬁcation [53]. There are lots of advantages of the technique over conventional molecular imprinting technique, as described in previous reports [54,55]. Only a few microliters of monomer solution are enough to polymerize dozens of samples at the same time in the same polymerization batch. Therefore, the method is useful to imprint templates which are very expensive or available in limited amounts. The technique also avoids potential solubility and conformational stability problems encountered with macromolecular targets, especially proteins. This is because immobilized templates are used in the process rather than adding the template to the monomer solution which otherwise is the standard procedure. This is also an important advantage for the ease of the template removal step after polymerization [54,55]. 6. Applications of Microcontact Imprinting Method with Capacitive Biosensors Ertürk et al. used microcontact imprinting method to prepare capacitive biosensors for various applications, please see Table 3. In their ﬁrst application, the authors used BSA as the model protein to prepare microcontact BSA imprinted capacitive biosensor [56]. In the ﬁrst step, the authors prepared glass cover slips with immobilized protein. This entity was called protein stamp. The cover slips were ﬁrst modiﬁed with 3-aminopropyl-triethoxysilane (APTES) to introduce amino groups on the surface and then with glutaraldehyde to modify these amino groups. Then, the cover slips were immersed in BSA solution overnight at 4 C. The electrode surface was on the other hand modiﬁed with polytyramine and then acryloyl chloride to introduce reactive groups on the surface which would be involved in the subsequent polymerization process. For the microcontact imprinting of BSA onto the gold electrode, the monomer solution containing methacrylic acid (MAA) as the functional monomer and poly-ethyleneglycoldimethacrylate (PEGDMA) as the cross-linker was prepared and the initiator was added into this solution. The modiﬁed gold electrode was treated with this monomer solution and then the protein stamp was brought into contact with this monomer treated electrode. Polymerization was initiated under UV light and ended in 15 min. After polymerization, the cover slip was removed from the surface with forceps and the microcontact BSA imprinted electrode was rinsed with water. The measurements were performed with a capacitive sensor involving the microcontact imprinted electrode inserted in an automated ﬂow injection system developed by Erlandsson et al. [57]. Sensors 2017, 17, 390 14 of 21 Table 3. Capacitive biosensors developed for different targets using microcontact imprinting method. Target Biosensing Method Monomers Dynamic Range LOD Selectivity Stability Ref. Capacitive biosensor Methacrylic acid (MAA); Poly Bovine Serum For human serum albumin >70 assays during 20 8 19 with current ethyleneglycol-dimethacrylate 1.0 10 M–1.0 10 M 1.0 10 M [56] Albumin (BSA) (HSA): 5%; For IgG: 3% 2 months pulse method (PEGDMA) Capacitive biosensor Prostate speciﬁc Selectivity coefﬁcient (k) = 2.27 About same level 17 10 17 with current MAA; EGDMA 2.0 10 M–2.0 10 M 16 10 M [58] antigen (PSA) for HSA, k = 2.02 for IgG during 50 injections pulse method HEMA; (2-Hydroxyethyl Capacitive biosensor K = 3.14 for B. subtilis, k = 3.32 methacrylate), About same level 2 7 1 1 E. coli with current 1.0 10 –1.0 10 CFUmL 70 CFUmL for S. aureus, k = 2.98 for [59] N-methacryloyl-L-histidine during 70 injections pulse method S. paratyphi methyl ester (MAH), EGDMA N-isopropylacrylamide (NIPAm), K = 733.1 for chymotrypsin Capacitive biosensor The loss in performance N,N-methylenebisacryl, amide (chy), k = 10.56 for BSA, 13 7 13 Trypsin with current 1.0 10 M–1.0 10 M 3.0 10 M was about 2% after [60] (MBAAm), Acrylamide, k = 6.50 for lysozyme (Lyz), pulse method 80 analyses Hydroxymethylacrylamide k = 3.46 for cytochrome c (cyt c) Sensors 2017, 17, 390 15 of 21 Sensors 2017, 17, x FOR PEER REVIEW 15 of 21 Capacitive biosensors based on a potential pulse have been used in many applications [1,11,61,62] Capacitive biosensors based on a potential pulse have been used in many applications including microcontact imprinting as seen in Table 3. In this type of capacitive biosensors, a small [1,11,61,62] including microcontact imprinting as seen in Table 3. In this type of capacitive biosensors, potential pulse is applied to the working electrode and the capacitance is measured. This potentiometric a small potential pulse is applied to the working electrode and the capacitance is measured. This pulse concept is sensitive to external electronic disturbances. Therefore, the system is prone to potentiometric pulse concept is sensitive to external electronic disturbances. Therefore, the system is inaccurate measurements and poor baseline stability. The sharp potential pulse may induce damage prone to inaccurate measurements and poor baseline stability. The sharp potential pulse may induce to the surface of the working electrode after a while which results in decreasing response of the damage to the surface of the working electrode after a while which results in decreasing response of electrode and eventually to replacement of the working electrode with a new electrode [57]. Another the electrode and eventually to replacement of the working electrode with a new electrode [57]. alternative way is to use a current pulse method to measure the capacitance at the electrode/solution Another alternative way is to use a current pulse method to measure the capacitance at the interface. Erlandsson et al. [57] developed a new concept to measure capacitance based on a constant electrode/solution interface. Erlandsson et al. [57] developed a new concept to measure capacitance current pulse to the biosensor transducer. In the basis of the principle, the system could be described based on a constant current pulse to the biosensor transducer. In the basis of the principle, the system as a simple resistor-capacitor (RC) circuit model. The schematic representation of the capacitance could be described as a simple resistor-capacitor (RC) circuit model. The schematic representation of measurement via current pulse method is shown in Figure 6. The system consists of: the capacitance measurement via current pulse method is shown in Figure 6. The system consists of: (1) A current source; (1) A current source; (2) An electro-chemical flow-cell which includes three electrodes: the working electrode which is a (2) An electro-chemical ﬂow-cell which includes three electrodes: the working electrode which is thin gold film coated with an insulating layer which functions as a bio-recognition layer to a thin gold film coated with an insulating layer which functions as a bio-recognition layer to immobilize the ligand, the auxiliary and reference electrodes which are made from a platinum immobilize the ligand, the auxiliary and reference electrodes which are made from a platinum wire; wire; (3) A potential differential ampliﬁer; and (3) A potential differential amplifier; and (4) A processor which converts the analogue potential to digital signal. (4) A processor which converts the analogue potential to digital signal. Figure 6. (a) Schematic representation of the capacitive system with current pulse method. The system Figure 6. (a) Schematic representation of the capacitive system with current pulse method. The system is comprised of: (1) current source; (2) flow cell which is connected to the working, reference and is comprised of: (1) current source; (2) ﬂow cell which is connected to the working, reference and auxiliary electrodes; (3) potential differential amplifier; and (4) a processor and ADC where the auxiliary electrodes; (3) potential differential ampliﬁer; and (4) a processor and ADC where the analogue potential is converted to digital signal; (b) A schematic view of Howland current pump used analogue potential is converted to digital signal; (b) A schematic view of Howland current pump used for supplying constant current; (c) Constant current supply to the sensor during the determined time for supplying constant current; (c) Constant current supply to the sensor during the determined time periods to measure the resistance and capacitance; (d) Capacitance is measured every minute and periods to measure the resistance and capacitance; (d) Capacitance is measured every minute and each minute (pulse) contains five sub pulse measurements with 20 ms intervals. (Reproduced from each minute (pulse) contains ﬁve sub pulse measurements with 20 ms intervals. (Reproduced from reference [57] with permission). reference [57] with permission). Sensors 2017, 17, 390 16 of 21 Sensors 2017, 17, 390 16 of 21 When When usin using g current pulse current pulse method, the method, the cap capacitive acitive measurement measurements s were wer perfo e performed rmed with with an an autautomated omated flow-inject ﬂow-injection ion syst system em as sho as shown wn in F in igFigur ure 7 [6 e 73] [63 . ]. Figure 7. Schematic representation of automated flow injection capacitive system. The components Figure 7. Schematic representation of automated ﬂow injection capacitive system. The components shown in the figure are integrated into a box to make a single, portable unit. (Reproduced from shown in the ﬁgure are integrated into a box to make a single, portable unit. (Reproduced from Reference [63] with permission). Reference [63] with permission). In their next study, Ertürk et al. [58] used microcontact imprinting method for detection of an In their next study, Ertürk et al. [58] used microcontact imprinting method for detection important biomarker, prostate specific antigen (PSA), for early detection of prostate cancer with of an important biomarker, prostate speciﬁc antigen (PSA), for early detection of prostate −17 capacitive biosensors. The standard solutions of different concentrations of PSA (2.0 × 10 M–2.0 × cancer with capacitive biosensors. The standard solutions of different concentrations of PSA −10 −15 10 M) were analysed by the system and the LOD value was calculated as 16 × 10 M. HSA and IgG 17 10 (2.0 10 M–2.0 10 M) were analysed by the system and the LOD value was calculated as were used as the competing proteins in order to test the selectivity of the system for PSA. These 16 10 M. HSA and IgG were used as the competing proteins in order to test the selectivity of −1 proteins are normally found in the levels of mg·mL where PSA is found at approximately the system for PSA. These proteins are normally found in the levels of mgmL where PSA is found −1 4.0 ng·mL in human serum. Therefore, when the developed system is tested against proteins at at approximately 4.0 ngmL in human serum. Therefore, when the developed system is tested concentrations of 1 mg/mL which are normally found in one-million-fold higher concentrations, the against proteins at concentrations of 1 mg/mL which are normally found in one-million-fold higher system showed around two times more selectivity for PSA compared to HSA and IgG. In the next concentrations, the system showed around two times more selectivity for PSA compared to HSA and step, the sensitivity and the selectivity of the MIP system were compared with the performance of the IgG. In the next step, the sensitivity and the selectivity of the MIP system were compared with the −14 system based on immobilized Anti-PSA antibody. The LOD was determined to be 12 × 10 M with performance of the system based on immobilized Anti-PSA antibody. The LOD was determined to be the Anti-PSA system. The results showed that the MIP capacitive system was very promising to 12 10 M with the Anti-PSA system. The results showed that the MIP capacitive system was very detect biomarkers which are important for diagnosis of various diseases. The authors compared the promising to detect biomarkers which are important for diagnosis of various diseases. The authors sensitivity of capacitive system for PSA detection with the microcontact imprinted surface plasmon compared the sensitivity of capacitive system for PSA detection with the microcontact imprinted −1 −14 resonance (SPR) system [64]. The LOD value was calculated around 91 pg·mL (18 × 10 M) with surface plasmon resonance (SPR) system [64]. The LOD value was calculated around 91 pgmL the SPR biosensors. This result proves that the capacitive system is approximately 1000 times more (18 10 M) with the SPR biosensors. This result proves that the capacitive system is approximately sensitive than the SPR system. 1000 times more sensitive than the SPR system. Microcontact imprinting method was used for E. coli detection by Idil et al. [59]. The authors Microcontact imprinting method was used for E. coli detection by Idil et al. [59]. The authors used a histidine containing specific monomer (N-methacryloyl-amido-histidine, MAH) as a metal used a histidine containing speciﬁc monomer (N-methacryloyl-amido-histidine, MAH) as a metal 2+ chelating ligand. By using a meal-chelate between MAH and copper(II) (MAH-Cu ), an enhanced 2+ chelating ligand. By using a meal-chelate between MAH and copper(II) (MAH-Cu ), an enhanced selectivity against certain amino acid residues present on the cell wall of E. coli was achieved. selectivity against certain amino acid residues present on the cell wall of E. coli was achieved. N-(hydroxyethyl)methacrylate (HEMA) was used as a functional monomer to make a complex with N-(hydroxyethyl)methacrylate (HEMA) was used as a functional monomer to make a complex 2+ 2+ MAH-Cu (pHEMA-MAH-Cu ) and EGDMA was used as cross-linker. The dynamic range for E. 2+ 2+ with MAH-Cu (pHEMA-MAH-Cu ) and EGDMA was used as cross-linker. The dynamic 2 7 −1 −1 coli detection was between 1.0 × 10 and 1.0 × 10 CFU·mL with a LOD value of 70 CFU·mL . Bacillus 2 7 1 range for E. coli detection was between 1.0 10 and 1.0 10 CFUmL with a LOD value of subtilis, Staphylococcus aureus and Salmonella paratyphi strains were used in selectivity experiments as 70 CFUmL . Bacillus subtilis, Staphylococcus aureus and Salmonella paratyphi strains were used in competing strains. The cross-reactivity ratios were between 24% and 58% against pre-mixed selectivity experiments as competing strains. The cross-reactivity ratios were between 24% and 58% suspensions of all bacterial strains. Even though the system showed cross-reactivity against against pre-mixed suspensions of all bacterial strains. Even though the system showed cross-reactivity competing strains, the ratio was negligible compared to the response of the system against E. coli. against competing strains, the ratio was negligible compared to the response of the system against River water and apple juice were used to show the detection performance of the system from E. coli. River water and apple juice were used to show the detection performance of the system from complex, real samples. Recovery value was found between 81% and 97% for E. coli detection from E. complex, real samples. Recovery value was found between 81% and 97% for E. coli detection from 2 4 −1 coli spiked (1.0 × 10 to 1.0 × 10 CFU·mL ) river water samples. The sensor had potential to monitor E. coli in contaminated water or food supplies. Sensors 2017, 17, 390 17 of 21 2 4 1 E. coli spiked (1.0 10 to 1.0 10 CFUmL ) river water samples. The sensor had potential to monitor E. coli in contaminated water or food supplies. Ertürk et al. [60] used microcontact imprinting method to develop capacitive biosensor for trypsin detection. A monomer solution containing hydroxymethylacrylamide, N-isopropylacrylamide (NIPAm), acrylamide and N,N-methylenebisacrylamide was prepared 0 0 and N,N,N ,N -tetramethylethylenediamine (TEMED, 5%, v/v) and ammonium persulfate (APS, 10%, v/v) were added into this solution. When the modiﬁed gold electrode was treated with monomer solution and Sensors br 2017 ought , 17, 390 into contact with the protein stamp which was carrying immobilized 17 of 21 trypsin on top, the polymerization continued for 3–5 h at room temperature. Schematic representation of Ertürk et al. [60] used microcontact imprinting method to develop capacitive biosensor for microcontact imprinting of trypsin on capacitive electrodes is shown in Figure 8. The dynamic range trypsin detection. A monomer solution containing hydroxymethylacrylamide, 13 7 13 for trypsin detection was between 1.0 10 M and 1.0 10 M with a LOD value of 3.0 10 M. N-isopropylacrylamide (NIPAm), acrylamide and N,N-methylenebisacrylamide was prepared and N,N,N′,N′-tetramethylethylenediamine (TEMED, 5%, v/v) and ammonium persulfate (APS, 10%, v/v) In order to test the selectivity and cross-reactivity of trypsin imprinted capacitive system for trypsin were added into this solution. When the modified gold electrode was treated with monomer solution (MW 23.3 kDa, isoelectric point (pI): 10.1–10.5), chymotrypsin (chy) (MW: 25.6 kDa, pI 8.3), bovine and brought into contact with the protein stamp which was carrying immobilized trypsin on top, the serum albumin (BSA) (MW: 66.5 kDa, pI 4.7), lysozyme (lyz) (MW 14.3 kDa, pI 11.35) and cytochrome c polymerization continued for 3–5 h at room temperature. Schematic representation of microcontact (cyt c) (MW: 12.3 kDa, pI 10.0–10.5) were selected as competing agents. The concentrations of the imprinting of trypsin on capacitive electrodes is shown in Figure 8. The dynamic range for trypsin −13 −7 −13 detection was between 1.0 × 10 M and 1.0 × 10 M with a LOD value of 3.0 × 10 M. In order to test proteins in the selectivity and cross-reactivity experiments were 1.0 mgmL . If it was tested with the the selectivity and cross-reactivity of trypsin imprinted capacitive system for trypsin (MW 23.3 kDa, 1 1 lower concentrations including 1.0 gmL or 1.0 pgmL , the selectivity results would be better isoelectric point (pI): 10.1–10.5), chymotrypsin (chy) (MW: 25.6 kDa, pI 8.3), bovine serum albumin because it would not be possible to detect interfering proteins in these concentration levels. Very low (BSA) (MW: 66.5 kDa, pI 4.7), lysozyme (lyz) (MW 14.3 kDa, pI 11.35) and cytochrome c (cyt c) (MW: afﬁnity of the system towards chy shows that the system can be used successfully for trypsin detection 12.3 kDa, pI 10.0–10.5) were selected as competing agents. The concentrations of the proteins in the −1 selectivity and cross-reactivity experiments were 1.0 mg·mL . If it was tested with the lower from pancreatic secretions where chy and try are found together. When the re-usability of the system −1 −1 concentrations including 1.0 µg·mL or 1.0 pg·mL , the selectivity results would be better because it was tested by monitoring the change in capacitance (-pFcm ) at the same concentration of trypsin would not be possible to detect interfering proteins in these concentration levels. Very low affinity of (10 M), the loss in the detection performance was around 2% after 80 analyses and there was not the system towards chy shows that the system can be used successfully for trypsin detection from any signiﬁcant difference in the performance after storage for two months at 4 C. In the last step, pancreatic secretions where chy and try are found together. When the re-usability of the system was −2 −7 tested by monitoring the change in capacitance (-pF·cm ) at the same concentration of trypsin (10 trypsin activity measured by capacitive system was compared with the trypsin activity measured by M), the loss in the detection performance was around 2% after 80 analyses and there was not any spectrophotometer at 410 nm. One unit of enzyme was deﬁned as the amount of enzyme catalyzing significant difference in the performance after storage for two months at 4 °C. In the last step, trypsin the conversion of one micromole of substrate, N -benzoyl-DL-arginine 4-nitroanilide hydrochloride activity measured by capacitive system was compared with the trypsin activity measured by (BAPNA), at 25 C, pH: 8.1 per minute. The trypsin activity measured spectrophotometrically was spectrophotometer at 410 nm. One unit of enzyme was defined as the amount of enzyme catalyzing 1 1 the conversion of one micromole of substrate, Nα-benzoyl-DL-arginine 4-nitroanilide hydrochloride 9 mUmL where the value was around 8.0 mUmL measured by capacitive system. The results (BAPNA), at 25 °C, pH: 8.1 per minute. The trypsin activity measured spectrophotometrically was 9 showed that there was correlation between two methods. The advantages of the developed method −1 −1 mU·mL where the value was around 8.0 mU·mL measured by capacitive system. The results including detection of trypsin in 20 min, high selectivity towards interfering proteins, high correlation showed that there was correlation between two methods. The advantages of the developed method with the spectrophotometer show that the system might be used successfully for detection of proteases including detection of trypsin in 20 min, high selectivity towards interfering proteins, high correlation with the spectrophotometer show that the system might be used successfully for detection and as a point of care system for diagnosis of pancreatic diseases. of proteases and as a point of care system for diagnosis of pancreatic diseases. Figure 8. Schematic representation of preparation of trypsin imprinted capacitive electrodes using Figure 8. Schematic representation of preparation of trypsin imprinted capacitive electrodes microcontact imprinting procedure: (A) preparation of glass cover slips (protein stamps); (B) using microcontact imprinting procedure: (A) preparation of glass cover slips (protein stamps); preparation of capacitive gold electrodes; and (C) imprinting of trypsin to the electrode surface via (B) preparation of capacitive gold electrodes; and (C) imprinting of trypsin to the electrode surface via microcontact imprinting method. (Reproduced from Reference [60] with permission). microcontact imprinting method. (Reproduced from Reference [60] with permission). Sensors 2017, 17, 390 18 of 21 7. Concluding Remarks From the above, it is obvious that the combination of capacitive biosensors and MIPs offers great possibilities, both with regard to sensitive and selective assays as well as to stable measurements of extended periods of time. 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This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (http://creativecommons.org/licenses/by/4.0/). http://www.deepdyve.com/assets/images/DeepDyve-Logo-lg.png Sensors (Basel, Switzerland) Pubmed Central http://www.deepdyve.com/lp/pubmed-central/capacitive-biosensors-and-molecularly-imprinted-electrodes-Lv2oYsPXSS

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Publisher: Pubmed Central
Copyright: © 2017 by the authors.
ISSN: 1424-8220
eISSN: 1424-8220
DOI: 10.3390/s17020390
Publisher site: See Article on Publisher Site

Abstract

sensors Review Capacitive Biosensors and Molecularly Imprinted Electrodes 1 , 1 , 2 Gizem Ertürk * and Bo Mattiasson CapSenze Biosystems AB, Lund 223 63, Sweden; [email protected] Department of Biotechnology, Lund University, Lund 222 40, Sweden * Correspondence: [email protected] Academic Editor: Nicole Jaffrezic-Renault Received: 14 December 2016; Accepted: 8 February 2017; Published: 17 February 2017 Abstract: Capacitive biosensors belong to the group of afﬁnity biosensors that operate by registering direct binding between the sensor surface and the target molecule. This type of biosensors measures the changes in dielectric properties and/or thickness of the dielectric layer at the electrolyte/electrode interface. Capacitive biosensors have so far been successfully used for detection of proteins, nucleotides, heavy metals, saccharides, small organic molecules and microbial cells. In recent years, the microcontact imprinting method has been used to create very sensitive and selective biorecognition cavities on surfaces of capacitive electrodes. This chapter summarizes the principle and different applications of capacitive biosensors with an emphasis on microcontact imprinting method with its recent capacitive biosensor applications. Keywords: capacitive biosensors; afﬁnity biosensors; microcontact imprinting 1. Introduction Afﬁnity biosensors can be divided into two main groups: those that measure direct binding between the target molecule and the afﬁnity surface on the sensor, and those biosensors which are adopted to binding assays using labelled reagents [1]. Biosensors operating with labelled afﬁnity-reagents are variations of conventional immunoassay technology in which ﬂuorescent markers, active enzymes, magnetic beads, radioactive species or quantum dots are generally used as labelling agents to label the target molecules [2,3]. Labelling is generally used to signiﬁcantly facilitate the signal generation and to conﬁrm the interaction between the probe and target molecules [4]. An important feature in using labelled reagents is the ampliﬁcation of the registered signals and thereby also the sensitivity one can reach. Assays based on the use of labelled reagents are time consuming and the labelled reagents are often expensive. Furthermore, assays with labelled reagents are usually multistep processes and that limit their application for real-time measurements. The design of label-free afﬁnity based biosensors is the objective of much current research. The aims are to establish alternative methods to the commercial ELISA-based immunoassays. The most attractive features of these types of biosensors are that they allow the monitoring of the analytes directly and in real-time [5]. Biosensors can be divided into four main groups according to transducer types [6]. These main groups involve: electrochemical transducers which involve potentiometric, voltammetric, conductometric, impedimetric and ﬁeld-effect transistors; optical transducers which include surface plasmon resonance (SPR) biosensors; piezoelectric transducers to which quartz crystal microbalance (QCM) biosensors can be given as an example; and thermometric transducers which measure the amount of heat with a sensitive thermistor to determine the analyte concentration. Sensors 2017, 17, 390; doi:10.3390/s17020390 www.mdpi.com/journal/sensors Sensors 2017, 17, 390 2 of 21 Among the different types of label-free biosensors, electrochemical biosensors have received particular attention owing to their properties [4]. These biosensors can also be miniaturized which is very important for many applications that need portable integrated systems. Miniaturization not only allows use at point of care, in a clinic, doctor ’s ofﬁce or at home but also reduces the cost of the diagnostic assays. Electrical biosensors fulﬁl these purposes as being fast, cheap, portable, miniaturized and label-free devices. Electrical biosensors can be classiﬁed as amperometric, voltametric impedance or capacitive sensors. This review deals with capacitive biosensors, different applications of capacitive biosensors developed for both detection of various targets and by using molecular imprinting technology with an emphasis on microcontact imprinting method. In this aspect, this is a novel review which includes the two technologies, capacitive biosensors and molecular imprinting; at the same time with lots of examples from published reports. The reports in molecular imprinting section were selected mainly from capacitive biosensors developed by using microcontact imprinting method. 2. Capacitive Biosensors Capacitive biosensors belong to the sub-category of impedance biosensors [3]. Capacitive biosensors measure the change in dielectric properties and/or thickness of the dielectric layer at the electrolyte-electrode interface when an analyte interacts with the receptor which is immobilized on the insulating dielectric layer [1]. The electric capacitance between the working electrode (an electrolytic capacitor/the ﬁrst plate) and the electrolyte (the second plate) is given by Equation (1) [7]: C = (" "A)/d (1) where " is the dielectric constant of the medium between plates, " is the permittivity of the free space 12 2 (8.85 10 F/m), A is the surface area of the plates (m ) and d is the thickness of the insulating layer (m). According to the given equation above, when the distance between the plates increases, the total capacitance decreases. In other words, in the assaying principle of this type of capacitive biosensors when a target molecule binds to the receptor, displacement of the counter ions around the capacitive electrode results in a decrease in the capacitance. The higher the amount of target molecules bound to the receptor is, the greater is the achieved displacement and the decrease in the registered capacitance [8]. The assaying principle of capacitive biosensors which are developed according to this rule is shown in Figure 1. The Equation (1) can be represented by two capacitors in series where the inner part includes the dielectric layer (C ) and the outer one corresponds to the biomolecule layer (C ). Then, the total dl bm capacitance (C ) can be described as Equation (2) [7]. 1 1 1 = + (2) C C C dl bm The electrochemical capacitors which are described based on the above-mentioned equation are known as constant phase element (CPE). The presence of CPE indicates that the observed capacitance of the system is frequency dependent. Capacitance can also be deﬁned as Equation (3) [7]. Z = (3) !C where Z is the impedance and ! is the radial frequency expressed in rads . This model implies that all of the measured current is capacitive. Sensors 2017, 17, 390 3 of 21 Sensors 2017, 17, 390 3 of 21 Figure 1. (A) Schematic diagram showing the change in capacitance (ΔC) as a function of time when Figure 1. (A) Schematic diagram showing the change in capacitance (DC) as a function of time when the analyte (IgG) interacts with the receptor molecule (Protein A) immobilized on the surface of the the analyte (IgG) interacts with the receptor molecule (Protein A) immobilized on the surface of the electrode. Subsequent rise in signal is due to the dissociation after the injection of the regeneration electrode. Subsequent rise in signal is due to the dissociation after the injection of the regeneration solution. In an ideal sensorgram, the baseline should turn back to the original level after regeneration solution. In an ideal sensorgram, the baseline should turn back to the original level after regeneration of of the surface; (B) Immobilization of the receptor molecule on the transducer surface via a the surface; (B) Immobilization of the receptor molecule on the transducer surface via a self-assembled self-assembled monolayer (SAM) of alkylthiols. When the target molecule interacts with the receptor, monolayer (SAM) of alkylthiols. When the target molecule interacts with the receptor, this creates this creates a double layer of counter ions around the gold transducer which results in a change in the a double layer of counter ions around the gold transducer which results in a change in the capacitance. capacitance. (Reproduced from Reference [8] with permission). (Reproduced from Reference [8] with permission). 3. Different Applications of Capacitive Biosensors 3. Different Applications of Capacitive Biosensors Different applications of capacitive biosensors developed for different targets are summarized Different applications of capacitive biosensors developed for different targets are summarized in in Table 1. Table 1. Sensors 2017, 17, 390 4 of 21 Table 1. Different applications of capacitive biosensors developed for different targets. Limit of Target Sensor Preparation Method Dynamic range (M) Selectivity Stability Ref. Detection (M) Immobilization of anti-CT antibodies on self-assembled monolayer (SAM) of 13 10 14 Cholera toxin (CT) 1.0 10 –1.0 10 1.0 10 uiu N/D N/D [9] lipoic acid and 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) Immobilization of anti-CT on Up to 36 times 18 12 20 Cholera toxin (CT) gold nanoparticles incorporated 0.1 10 –10 10 9.0 10 N/D [10] with an RSD of 2.5% on a poly-tyramine layer Proteins Immobilization of anti-HIV 1 p24 antigen on gold nanoparticles 20 17 20 HIV-1 p24 antigen 10.1 10 –10.1 10 3.32 10 N/D N/D [11] incorporated on a poly-tyramine layer Immobilization of anti-VEGF aptamer ﬁrst capturing the VEGF 14 11 VEGF protein then, sandwiching with N/D N/D N/D [12] 13 10 –2.6 10 antibody-conjugated magnetic beads Oligo-T was used as the competing agent, when the Covalent attachment of 25-mer temperature was increased 8 11 11 25-mer oligo C oligo C on poly-tyramine 10 –10 10 N/D [13] from RT to 50 C, the DC modiﬁed electrode value decreased from 2 2 48 nFcm to 3 nFcm GOPTS functionalized Thiol modiﬁed oligonucleotides surfaces were more stable at were immobilized on Au and 4 C. Ten-fold decrease in 6 3 ssDNA N/D N/D [14] 0.5 10 –1.0 10 Nucleic acids 3-glycidoxypropyl-tri-methoxy ﬂuorescence intensity after silane (GOPTS) 1 week even when the substrates were stored at 4 C. Immobilization of ssDNA 10 11 Nampt aptamers on SAM of 0–45 10 1.8 10 N/D N/D [15] mercaptopropionic acid (MPA) Complementary DNA Immobilization of pyrrolidinyl Could be reused for provided a much higher DC 11 10 12 Target DNA peptide nucleic acid probes 1.0 10 –1.0 10 6–10 10 58–73 times with an average [16] compared to single and (acpcPNA) residual activity of 98% double mismatched DNA Sensors 2017, 17, 390 5 of 21 Table 1. Cont. Limit of Target Sensor Preparation Method Dynamic range (M) Selectivity Stability Ref. Detection (M) Based on the interaction For the ﬁrst 35 cycles, the between E. coli and concanavalin residual activity was 1 6 1 1 Total bacteria 12 CFUmL –1.2 10 CFUmL 12 CFUmL N/D [17] A immobilized on a modiﬁed 95% 3% (RSD = 3.2%). Cells gold surface After 35 cycles, it was 85%. E. coli cells immobilized on SAM 5 1 7 1 E. coli N/D N/D N/D [18] 8 10 CFUmL –8 10 CFUmL of Mercaptopropionic acid (MPA) Immobilization of metal Hg(II), Cu(II), 15 3 resistance and metal regulatory 10 –10 N/D N/D N/D [19] Zn(II), Cd(II) proteins on gold electrode 1. Immobilization of whole bacterial cell to emit Heavy metals 84% of the activity loss a bioluminescent/ﬂuorescent 0–200 10 within 6 days N/D [20] 1.0 10 Cu(II), Cd(II), signal in the presence of heavy Hg(II) metal ions 2. Immobilization of heavy metal 15 1 10 –10 Stable over 16 days binding proteins Small sugars including D-fructose, D-mannose, Immobilization of ConA on gold A neglectable loss in D-maltose, 6 2 6 Glucose nanoparticles incorporated on the 1.0 10 –1.0 10 1.0 10 sensitivity after [21] Saccharides methyl--D-glucopyranoside, tyramine modiﬁed gold electrode 10 cycles (7.5%) methyl--D-mannopyranoside also bound instead of glucose Immobilization of ConA and Small molecules and high 5 1 6 Glucose replacement of small glucose 1.0 10 –1.0 10 1.0 10 molecular weight dextran N/D [22] with the large glucose polymer also bound instead of glucose Immobilization of molecularly Cross reactant contribution 6 6 Metergoline imprinted spherical beads on 1–50 10 1.0 10 N/D [23] was maximum 1.3 nF modiﬁed gold electrode Competing agents’ binding Little variation over 28 6 9 12 Aﬂatoxin B1 Bioimprinting 3.2 10 –3.2 10 6.0 10 was signiﬁcantly lower than injections with non-reduced [24] Small molecules aﬂatoxin B1 Schiff’s bases Monoclonal anti-OTA Differences for ochratoxin B Ochratoxin A immobilization on Si N 3 4 12 12 2.47–49.52 10 4.57 10 and aﬂatoxin B1 were N/D [25] (OTA) substrate combined with not signiﬁcant magnetic nanoparticles (MNPs) Sensors 2017, 17, 390 6 of 21 3.1. Protein Detection Labib et al. [9] developed a sensitive method for detection of cholera toxin (CT) using a ﬂow-injection capacitive immunosensor based on self-assembled monolayers. Monoclonal antibodies against the subunit of CT (anti-CT) were immobilized on the gold electrode surface which was modiﬁed by lipoic acid and 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC). The immunosensor showed linear response to CT concentrations in the concentration range between 1.0 10 M and 10 14 1.0 10 M under optimized conditions. Limit of detection (LOD) value was 1.0 10 M. The LOD value obtained from capacitive immunosensor was compared with the LOD values that were obtained from sandwich ELISA and SPR based immunosensors. The ELISA had a LOD of 12 11 1.2 10 M whereas SPR had a LOD value of 1.0 10 M. The results proved that the method is more sensitive than the other two techniques used in this study. In another study by the same research group, a label-free capacitive immunosensor was developed for direct detection of CT present at sub-attomolar level. Gold nanoparticles (AuNPs) were incorporated on a polytyramine modiﬁed gold electrode and anti-CT antibody was immobilized on this surface. Tyramine provides free amino groups which are very useful to immobilize the afﬁnity ligand to the transducer. At the same time very thin and uniform ﬁlms can be formed on the electrode surfaces during the electro-polymerization of tyramine. After the immobilization step, the formation of antigen-antibody complexes resulted in a change in capacitance and by this way the concentration of CT was determined. The dynamic range was between 0.1 aM and 10 pM where the LOD value was 9 10 M (0.09 aM). The electrode could be regenerated with a good reproducibility for up to 36 times with a relative standard deviation (RSD) value of 2.5%. Real sample analyses were performed from water samples collected from a local stream and matrix effect was eliminated with a 10.000 times dilution prior to analysis. The developed system had potential to be used as a portable electrochemical analyser for ﬁeld conditions [10]. The same strategy was used to develop a capacitive biosensor for sensitive detection of HIV-1 p24 antigen [11]. Following polytyramine electro-polymerization on the gold electrode surface, gold nanoparticles were incorporated onto the electrode and then, anti-HIV-1 p24 monoclonal antibodies were immobilized on top. HIV-1 p24 antigen was detected from standard p24 solutions in the 20 17 concentration range of 10.1 10 to 10.1 10 M. The reasons for extreme sensitivity of the capacitive biosensors were explained by the authors in two ways. Firstly, the capacitive technique is very convenient to detect the size and position of the electrical double layer formed at the interface. However, in other electrochemical techniques such as the amperometric technique, the detection is based on the measurement of only the transport of electrons. The second factor is the increased surface area on the electrode via the use of immobilized AuNPs. The gold nanoparticles also signiﬁcantly contributed to the decrease in the interfacial resistance which facilitated the electron transfer at the electrode surface [11]. Qureshi et al. [12] developed a capacitive aptamer based sensor for detection of vascular endothelial growth factor (VEGF) in human serum. Systematic evolution of ligands by exponential enrichment (SELEX) process was utilized to select the highly speciﬁc and selective anti-VEGF aptamer to bind VEGF to the aptasensor surface. When a sandwich assay was tested by forming sandwich complex with anti-VEGF aptamer+VEGF+anti-VEGF antibody, the generated signal was enhanced by 3–8 folds compared to the direct assay. The developed sensor showed a dynamic detection range 14 11 from 13 10 M to 2.6 10 M of VEGF protein in human serum. The results showed that the developed system could be successfully used in clinical diagnosis to detect biomarkers in real samples in a convenient and sensitive way. 3.2. Nucleic Acid Detection Mahadhy et al. [26] used in a model study the capacitive sensor to monitor the capture of complementary single-stranded nucleic acids. A 25-mer oligo-C was immobilized onto the polytyramine modiﬁed gold electrode surface. Temperature was raised to 50 C to reduce non-speciﬁc Sensors 2017, 17, 390 7 of 21 hybridization in order to increase the selectivity and hybridization was used in order to amplify the signal by using longer nucleic acid molecules. Later, Mahadhy et al. [13] developed a promising ultrasensitive, automated ﬂow-based and portable gene sensor. The PCR-free biosensor proved the possibility for powerful detection of foodborne pathogens in diagnostic situations and multi-drug resistant bacteria in the near future. Rapid detection of foodborne pathogens is crucial before many people are infected while early detection of multi-drug resistant bacteria is important to isolate the infected patients earlier and to reduce the risk of spreading. Two functionalization layers; gold (Au) and 3-glycidoxypropyl-tri-methoxy silane (GOPTS) were used to immobilize thiol modiﬁed oligonucleotides on silicon surfaces. GOPTS showed better performance as a functionalization layer because the hybridization efﬁciency was higher, the stability over time was better and regeneration of the surface after analyte binding was easier. Therefore for the development of microcantilever or micro-membrane based biosensors, GOPTS might be a more promising alternative [14]. A single stranded DNA (ssDNA) aptamer was developed to bind to nicotinamide phosphoribosyl transferase (Nampt) through SELEX and implemented in a capacitive biosensor [15]. The LOD for 11 10 Nampt was 1.8 10 M with a dynamic detection range in serum of up to 9 10 M. Nampt is an important biomarker for obesity-related metabolic diseases, some types of cancers and chronic 1 11 diseases. Normal level of Nampt in human plasma is around 15 ngmL (27 10 M). Therefore, the developed system has potential as a diagnostic tool and in point-of-care applications. Pyrrolidinyl peptide nucleic acid probes were immobilized onto the self-assembled monolayer (SAM) modiﬁed gold electrode surface to develop a DNA capacitive biosensor [16]. Four different alkanethiols with various chain lengths were used as a SAM to determine the inﬂuence of the length and the terminating head group of blocking thiols on the sensitivity and speciﬁcity. In the study, the blocking thiol which had an equal length to the –OH terminating head group gave the highest sensitivity and high binding speciﬁcity. Thus far, there are no good examples on use of MIPs in connection to monitoring of speciﬁc DNA sequences. Since the potential is large, one can expect such developments to happen in the near future. 3.3. Cell Detection Jantra et al. [17] developed a label-free afﬁnity biosensor for detecting and enumerating total bacteria based on the interaction between E. coli and Concanavalin A (Con A) immobilized on a modiﬁed gold surface. The analyses were completed in less than 20 min with both SPR and 7 1 impedimetric capacitive biosensor. Compared to SPR (LOD: 6.1 10 CFUmL ), the capacitive system showed much higher sensitivity (LOD: 12 CFUmL ). The developed system might be used successfully for total bacteria analysis from water sources. Rydosz et al. [27] developed a new type of label-free microwave sensor in a form of interdigitated capacitor for bacterial lipopolysaccharide detection. The sensor surface was coated with T4 phage gp37 adhesin. The adhesin molecule bound E. coli by recognizing its bacterial host lipopolysaccharide (LPS). The binding was highly speciﬁc and irreversible. Recognition between the phage adhesion and bacterial LPS was based on the recognition of saccharide determinants of LPS which means very speciﬁc determination of bacterial strain or its endotoxins within the genus and the species. The selectivity experiments showed that the response for speciﬁc LPSs was signiﬁcantly different from the reference measurements and the response for the non-speciﬁc LPSs was very close to the reference values. The developed method was promising for label-free LPS detection and could be used as an alternative for ﬁber-optic, electrochemical and classic biochemical and immunochemical methods. In their other study [28], same authors used bacteriophage-adhesin-coated long-period gratings for recognition of bacterial lipopolysaccharides. Long-period gratings (LPG) bio-functionalization methodology was based on coating the LPG surface with nickel ions which were capable of binding of gp37-histidine tag. The advantage of using adhesins for the bio-functionalization of the biosensor was to give ability Sensors 2017, 17, 390 8 of 21 for low-cost bio-sensitive molecule exchange and surface regeneration. In this work, for the ﬁrst time, adhesion has been applied for bacteria and their endotoxin detection. T4 phage adhesin bound E. coli B LPS in its native or denatured form in a highly speciﬁc and irreversible way. Rocha et al. [29] used alternating current electrokinetics (ACEK) capacitive sensing to detect and quantify the microbial cell abundance in aquatic systems. Microbial abundance was detected by measuring the electrical signal. Three different microbial cell cultures including Bacillus subtilis, Alcanivorax borkumensis and Microcystis aeruginosa were detected by using the developed system. The results showed that the sensor is capable of reliably detecting microbial cells even though they have major physiological differences between Gram-positive (B. subtilis), Gram-negative (A. borkumensis) and cyanobacteria (M. aeruginosa). The system is promising to detect and estimate microorganism population sizes in batch cultures, environmentally sourced seawater and groundwater systems. For expanded use and commercialization of nanotechnology products, toxicity determination is an important ﬁeld application. For this purpose, Qureshi et al. [18] developed a whole-cell based capacitive biosensor to determine the biological toxicity of nanoparticles (NPs). They used iron oxide (Fe O ) nanoparticles as models in the study. The living E. coli cells were immobilized on the capacitive 3 4 sensor chips. Then, these chips were interacted with different sizes of Fe O NPs (5, 20 and 100 nm). 3 4 The smallest Fe O NPs resulted in a maximum capacitance change because they were able to interact 3 4 with E. coli cells on the sensor chip very efﬁciently. The morphological changes on the surface of E. coli cells after interacting with Fe O NPs were examined with SEM. 3 4 3.4. Heavy Metal Detection Two metal binding proteins were over-expressed in E. coli, puriﬁed and immobilized on a thiol-modiﬁed capacitive sensor surface. Capacitive sensor was used to monitor conformational changes following heavy metal binding including copper, cadmium, mercury and zinc. Metal ion detection could be done in down to femtomolar concentrations with the developed system [30]. The same group expressed and puriﬁed metal resistance and metal regulatory proteins from bacterial strains and immobilized these proteins on the capacitive biosensor surface for heavy metal detection [19]. The system allowed the detection of heavy metals including Hg(II), Cu(II), Zn(II) and Cd(II) in pure solutions down to 10 M concentrations. Corbisier at al. [20] established two different biosensor technologies for detection of several heavy metal ions in environmental samples. The principle of the ﬁrst approach was to develop whole cell bacterial biosensors which emitted a bioluminescent/ﬂuorescent signal in the presence of heavy metal ions. In the second approach, direct interaction between metal binding proteins and heavy metal ions was used as the detection principle in the capacitive biosensors. In the study, the main advantage of the whole cell based sensors was their ability to react only to biologically available metal ions, whereas the latter one (protein based sensors) was more sensitive towards metal ions. MIPs selective for heavy metal ions have been presented in connection to separation technology [31–36]. It is an obvious development that also sensors for heavy metal ions will be developed based on selective MIPs. 3.5. Saccharide Detection By using capacitive biosensors, Labib et al. [21] used the capacitive biosensor to detect glucose based on gold nanoparticles which were ﬁxed on a poly-tyramine modiﬁed gold electrode surface. Dextran (MW: 39 kDa) was used as a regeneration agent by utilizing a competitive assay for glucose in 6 2 the study. The dynamic range for glucose detection was between 1.0 10 and 1.0 10 M with a LOD value of 1.0 10 M. By using capacitive biosensors, the same authors [22] developed a technique based on the competition between a small molecular mass analyte and a large analyte-carrier conjugate. In the basis of the technique, when a large glucose polymer binds to the biorecognition molecule (Con A) immobilized on the electrode surface, it results in a decrease in the capacitance. Then, in the Sensors 2017, 17, 390 9 of 21 next step, when the low molecular mass analyte (glucose) is introduced to the system, the effect is reverse, the small glucose molecule will replace the large glucose polymer which is bound on the immobilized Con A as shown in Figure 2. By measuring the shift-back in capacitance, the glucose concentration could be determined by the technique. The authors used this technique to measure IgG as a glycoconjugate and detect its aggregation using immobilized Con A. When the glycoconjugate (IgG) was injected, the decrease in capacitance was measured to determine its concentration. In the second step, when concentrated glucose was injected into the system, the increase in capacitance was employed to determine the glucose concentration. The results showed that this technique is promising for monitoring Sensors 2017, 17 small , 390 molecules with high sensitivity and broad detection range. 10 of 21 Figure 2. Schematic representation of the competitive glucose binding assay. (a) When glucose is Figure 2. Schematic representation of the competitive glucose binding assay. (a) When glucose is injected into the capacitive system, it binds to the immobilized Concanavalin A (ConA) on the surface. injected into the capacitive system, it binds to the immobilized Concanavalin A (ConA) on the surface. However, this binding does not make any change in the capacitance level, as shown in the graph on However, this binding does not make any change in the capacitance level, as shown in the graph on the the right, due to the small size of the glucose molecule; (b) When a glucose polymer (dextran) is right, due to the small size of the glucose molecule; (b) When a glucose polymer (dextran) is injected injected into the system, binding of this big polymer to ConA results in a decrease in the capacitance into the system, binding of this big polymer to ConA results in a decrease in the capacitance signal; signal; (c) When glucose is injected into the system again, displacement of dextran with glucose results (c) When glucose is injected into the system again, displacement of dextran with glucose results in the in the capacitance turn back to the original baseline level. (Reproduced from Reference [22] with capacitance turn back to the original baseline level. (Reproduced from Reference [22] with permission). permission). 3.6. Small Organic Molecules The radio frequency (RF) detection method is one of the promising methods for glucose detection out of several detection methods available [37]. When an analyte is injected into the RF biosensor, Small molecules such as pesticides, herbicides, and antibiotics are widely discarded and encountered in naturally flowing waters. These pollutants in environment have an impact on changes will occur owing to the inductive and capacitive effects [38]. These changes will cause communities and eco-systems. Therefore, detection of these molecules in a sensitive, cheap, robust losses and considerable shifts in the resonance frequency of the device. The change in capacitance is and fast way is crucial. Lenain et al. [23] chose metergoline as a model compound representingsmall proportional to the dielectric constant and the distance between the biomolecule layer and the dielectric organic molecules such as pharmaceutical residues. Emulsion polymerization was used to produce layer. A reusable robust RF biosensor was developed by Kim et al. [38] to monitor real-time glucose small, uniformly sized, spherical MIPs. These MIP beads were attached to the poly-tyramine level in human serum. The resonance behaviour of the system was analysed with human serum modified gold surface. Scanning electron microscope (SEM) images of the electrode surface are 1 1 samples containing different glucose concentrations ranging from 148–268 mgdL , 105–225 mgdL −6 −6 shown in Figure 3. Working range for metergoline was from 1.0 × 10 M to 50 × 10 M with a LOD and at a deionized water glucose concentration in the range of 25–500 mgdL . The response time for −6 value of 1.0 × 10 M. In cross reactivity analysis, even though the structural analogs showed binding, glucose was measured as 60 s with a LOD value of 8.01 mgdL . A total of 21 different experiments this contribution was only around 1.3 nF. The sensor response was more stable at higher ionic for each strength but the extent of concentration of serum ca and paci D-gl tance cha ucosensolution ge for different co were analysed ncentrations for reusability of analyte w and the as less relative pronounced compared to lower electrolyte concentrations. standard deviation (RSD) was less than 1% for each concentrations of serum samples and aqueous Bioimprinting is a technology used to mimic specific sites for modification of biological D-glucose solutions. molecules. The process consists of four steps, as shown in Figure 4 [24]. In the area of MIPs used for bioseparation, much work has been done concerning carbohydrates (1) Unfold [39 ing – t 41 he conform ]. It is obvious ation of t that he st one arting prot can make ein un good der acid MIPs ic con with ditions high ; efﬁciency in binding (2) Addition of template molecule and allow interaction between the template molecule and the denatured protein in order to form new molecular configurations; (3) Cross-linking of the protein to stabilize the new molecular protein conformation; and (4) Dialysis to remove the template molecule. Sensors 2017, 17, 390 10 of 21 target molecules or fragments thereof. Based on the observations from afﬁnity chromatography, one can foresee that such systems will also soon be presented for capacitive biosensors [42–48] (Table 2). Table 2. Molecularly imprinted polymers (MIPs) produced with high binding efﬁciency for afﬁnity chromatography applications. Template Method Matrix Comments Ref. Preconcentration on BAP BAP-imprinted poly with HPLC equipped with Benzo[a]pyrene (2-hydroxyethylmethacrylate- Aqueous solutions [42] a ﬂuorescence detector (BAP) N-methacryloyl-(L)-phenylalanine (HPLC-FLD) composite cryogel cartridge MIP-solid phase extraction Melamine imprinted Extraction and enrichment Melamine Water + milk [43] monolithic cartridges of melamine Cholesterol imprinted Gastrointestinal Cholesterol adsorption Cholesterol [44] polymeric nanospheres mimicking solution Iron chelated poly Catalase puriﬁcation (2-hydroxyethylmethacrylate- Catalase Rat liver [45] from rat liver N-methacryloyl-(L)-glutamic acid cryogel discs Chiral separation of l-phenylalanine with L-phenylalanine L-Phe imprinted Aqueous solutions [46] FPLC (fast protein (L-Phe) cryogel cartridges liquid chromatography) Separation of triazine with Triazine imprinted capillary electro- Triazine Aqueous solutions [47] monolithic columns chromatography (CEC) Cytochrome c puriﬁcation Surface imprinted bacterial Cytochrome c Rat liver [48] from rat liver cellulose nanoﬁbers 3.6. Small Organic Molecules Small molecules such as pesticides, herbicides, and antibiotics are widely discarded and encountered in naturally ﬂowing waters. These pollutants in environment have an impact on communities and eco-systems. Therefore, detection of these molecules in a sensitive, cheap, robust and fast way is crucial. Lenain et al. [23] chose metergoline as a model compound representingsmall organic molecules such as pharmaceutical residues. Emulsion polymerization was used to produce small, uniformly sized, spherical MIPs. These MIP beads were attached to the poly-tyramine modiﬁed gold surface. Scanning electron microscope (SEM) images of the electrode surface are shown in Figure 3. 6 6 6 Working range for metergoline was from 1.0 10 M to 50 10 M with a LOD value of 1.0 10 M. In cross reactivity analysis, even though the structural analogs showed binding, this contribution was only around 1.3 nF. The sensor response was more stable at higher ionic strength but the extent of capacitance change for different concentrations of analyte was less pronounced compared to lower electrolyte concentrations. Bioimprinting is a technology used to mimic speciﬁc sites for modiﬁcation of biological molecules. The process consists of four steps, as shown in Figure 4 [24]. (1) Unfolding the conformation of the starting protein under acidic conditions; (2) Addition of template molecule and allow interaction between the template molecule and the denatured protein in order to form new molecular conﬁgurations; Sensors 2017, 17, 390 11 of 21 (3) Cross-linking of the protein to stabilize the new molecular protein conformation; and (4) Dialysis to remove the template molecule. Sensors 2017, 17, 390 11 of 21 Sensors 2017, 17, 390 11 of 21 c d c d Figure 3. Scanning electron microscope (SEM) pictures of the electrode surface after functionalization Figure 3. Scanning electron microscope (SEM) pictures of the electrode surface after functionalization Figure 3. Scanning electron microscope (SEM) pictures of the electrode surface after functionalization with imprinted polymers. From left to right, top to bottom: (a) SEM picture of electrode surface; (b,c) with imprinted polymers. From left to right, top to bottom: (a) SEM picture of electrode surface; with imprinted polymers. From left to right, top to bottom: (a) SEM picture of electrode surface; (b,c) SEM pictures of centre of the electrode; and (d) SEM picture of the border between the gold layer and (b,c) SEM pictures of centre of the electrode; and (d) SEM picture of the border between the gold layer SEM pictures of centre of the electrode; and (d) SEM picture of the border between the gold layer and wafer. (Reproduced from Reference [23] with permission). and wafer. (Reproduced from Reference [23] with permission). wafer. (Reproduced from Reference [23] with permission). Figure 4. Schematic representation of bio-imprinting process. (Reproduced from Reference [24] with Figure 4. Schematic representation of bio-imprinting process. (Reproduced from Reference [24] with permission). permission). Figure 4. Schematic representation of bio-imprinting process. (Reproduced from Reference [24] with permission). Gutierrez et al. [24] used bioimprinting to develop a capacitive biosensor for aflatoxin detection. Gutierrez et al. [24] used bioimprinting to develop a capacitive biosensor for aﬂatoxin detection. Aflatoxins are natural food contaminants with a high risk for human health. Ovalbumin was used as Gutierrez et al. [24] used bioimprinting to develop a capacitive biosensor for aflatoxin detection. Aﬂatoxins platform ar fo e natural r bioimprintin food contaminants g of aflatoxin because with a when bovine serum albumi high risk for human health. n (BSA) w Ovalbumin as used, there was used Aflatoxins are natural food contaminants with a high risk for human health. Ovalbumin was used as was no change in capacitance owing to the high hydrophobicity of sites of BSA. Three competitive as platform for bioimprinting of aﬂatoxin because when bovine serum albumin (BSA) was used, platform mycotoxi for bioimprintin ns were used in the cross-rea g of aflatoxin because ctivity a when bovine serum albumi nalysis and the changes inn ca (BSA) w pacitaance were s used, there there was no change in capacitance owing to the high hydrophobicity of sites of BSA. Three competitive significantly lower than that registered from aflatoxin solution. was no change in capacitance owing to the high hydrophobicity of sites of BSA. Three competitive mycotoxins were used in the cross-reactivity analysis and the changes in capacitance were signiﬁcantly mycotoxins were used in the cross-reactivity analysis and the changes in capacitance were lower than that registered from aﬂatoxin solution. significantly lower than that registered from aflatoxin solution. Sensors 2017, 17, 390 12 of 21 Sensors 2017, 17, 390 12 of 21 Silicon nitride substrate (Si N ) combined with magnetic nanoparticles (MNPs) was used to 3 4 develop a capacitive immunosensor for ochratoxin A (OTA) detection. Silicon nitride allows an easy Silicon nitride substrate (Si3N4) combined with magnetic nanoparticles (MNPs) was used to control of the ﬁlm composition and thickness and also prevents the undesirable impurities. These are develop a capacitive immunosensor for ochratoxin A (OTA) detection. Silicon nitride allows an easy the main advantages of the substrate used in the study. Magnetic nanoparticles comprised of control of the film composition and thickness and also prevents the undesirable impurities. These are a conductive core and a carboxylic acid modiﬁed shell which was used to immobilize OTA antibodies. the main advantages of the substrate used in the study. Magnetic nanoparticles comprised of a The LOD value was calculated as 4.57 10 M in the study and the selectivity results against conductive core and a carboxylic acid modified shell which was used to immobilize OTA antibodies. ochratoxin B and aﬂatoxin G1 showed that −the 12 potential difference was not so signiﬁcant when The LOD value was calculated as 4.57 × 10 M in the study and the selectivity results against compared to the difference for OTA detection [25]. ochratoxin B and aflatoxin G1 showed that the potential difference was not so significant when compared to the difference for OTA detection [25]. 4. Molecular Imprinting 4. Molecular Imprinting During the early 1970s, Wulff et al. [49] and Klotz et al. [50] introduced the molecular imprinting to imprint templates in organic polymers. Then, Mosbach et al. [51] reported the use of molecularly During the early 1970s, Wulff et al. [49] and Klotz et al. [50] introduced the molecular imprinting imprinted polymers (MIPs) in biosensors instead of antibodies which was a breakthrough. to imprint templates in organic polymers. Then, Mosbach et al. [51] reported the use of molecularly imprint The formation ed polymers of ( MIPs MIPs) involves in biosensor three s in steps: stead of antibodies which was a breakthrough. The formation of MIPs involves three steps: (1) Pre-complexation of functional monomers around the template molecule in solution either by (1) Pre-complexation of functional monomers around the template molecule in solution either by forming covalent bonds or by self-assembling with non-covalent bonds; forming covalent bonds or by self-assembling with non-covalent bonds; (2) Polymerization of the resulting complex in the presence of cross-linking monomers and suitable (2) Polymerization of the resulting complex in the presence of cross-linking monomers and suitable solvents/ionic liquids as porogens; and solvents/ionic liquids as porogens; and (3) Removal of template molecule from the synthesized polymer. (3) Removal of template molecule from the synthesized polymer. The The resu resulting lting MIP MIP cont contains ainsr recogn ecognition ition cavities cavities c capable apable of ofsel selective ective recogn recognition ition of compound of compounds s that fit these cavities with respect to shape, size, position and orientation of the recognition sites [52]. that ﬁt these cavities with respect to shape, size, position and orientation of the recognition sites [52]. How MIPs can mimic natural recognition units in different applications are shown schematically in How MIPs can mimic natural recognition units in different applications are shown schematically in Figure 5. Figure 5. Figure 5. Different applications of MIPs in: (A) immunosensors; (B) enzyme-linked immunosorbent Figure 5. Different applications of MIPs in: (A) immunosensors; (B) enzyme-linked immunosorbent assay (ELISA); (C) enzyme electrodes, reaction rate and analyte concentration of enzyme electrodes assay (ELISA); (C) enzyme electrodes, reaction rate and analyte concentration of enzyme and catalytic MIP-coated electrodes can be estimated by electroactive substrate/product electrodes and catalytic MIP-coated electrodes can be estimated by electroactive substrate/product consumption/production during the catalytic reaction or electron transfer from the electrode surface consumption/production during the catalytic reaction or electron transfer from the electrode surface to to the active centre of enzyme/MIP; (D) DNA chips; and (E) enzyme immobilization and competitive the active centre of enzyme/MIP; (D) DNA chips; and (E) enzyme immobilization and competitive binding of the analyte. (Reproduced from Reference [52] with permission). binding of the analyte. (Reproduced from Reference [52] with permission). MIP technology has successfully been used for imprinting of low molecular weight templates. However there are still some difﬁculties of molecular imprinting technique when it is used for Sensors 2017, 17, 390 13 of 21 macromolecular templates including proteins. Due to this, many researchers have focused on the alternative techniques including imprinting the template directly onto a substrate or immobilizing the template protein on a glass support and use it as a protein stamp. The latter is called microcontact imprinting. 5. Microcontact Imprinting Microcontact imprinting technique was ﬁrst introduced by Chou et al. [53]. In the study, the authors formed the microcontact imprints between two cleaned glass surfaces. Template protein was immobilized on the cover slip and then, functional monomer was added on top in order to allow site-speciﬁc organization of the functional monomer by the template. In the next step, a drop of solution including cross-linker and initiator was dropped on the pre-modiﬁed support glass. Support glass was brought into contact with the cover slip to provide the two functionalized surfaces to contact. The glass assemblies were placed in a UV reactor to initiate the polymerization and it continued for 17 h. After polymerization, the cover slip was removed from the surface with forceps and the support was washed with various solutions. Polymer ﬁlm thickness measurements showed it to be around 10 m. The imprinted polymers showed better selectivity for their native templates than competing proteins. By this study, it was reported that microcontact imprinting technique might be used successfully for relatively large proteins in biosensor applications for detection and quantiﬁcation [53]. There are lots of advantages of the technique over conventional molecular imprinting technique, as described in previous reports [54,55]. Only a few microliters of monomer solution are enough to polymerize dozens of samples at the same time in the same polymerization batch. Therefore, the method is useful to imprint templates which are very expensive or available in limited amounts. The technique also avoids potential solubility and conformational stability problems encountered with macromolecular targets, especially proteins. This is because immobilized templates are used in the process rather than adding the template to the monomer solution which otherwise is the standard procedure. This is also an important advantage for the ease of the template removal step after polymerization [54,55]. 6. Applications of Microcontact Imprinting Method with Capacitive Biosensors Ertürk et al. used microcontact imprinting method to prepare capacitive biosensors for various applications, please see Table 3. In their ﬁrst application, the authors used BSA as the model protein to prepare microcontact BSA imprinted capacitive biosensor [56]. In the ﬁrst step, the authors prepared glass cover slips with immobilized protein. This entity was called protein stamp. The cover slips were ﬁrst modiﬁed with 3-aminopropyl-triethoxysilane (APTES) to introduce amino groups on the surface and then with glutaraldehyde to modify these amino groups. Then, the cover slips were immersed in BSA solution overnight at 4 C. The electrode surface was on the other hand modiﬁed with polytyramine and then acryloyl chloride to introduce reactive groups on the surface which would be involved in the subsequent polymerization process. For the microcontact imprinting of BSA onto the gold electrode, the monomer solution containing methacrylic acid (MAA) as the functional monomer and poly-ethyleneglycoldimethacrylate (PEGDMA) as the cross-linker was prepared and the initiator was added into this solution. The modiﬁed gold electrode was treated with this monomer solution and then the protein stamp was brought into contact with this monomer treated electrode. Polymerization was initiated under UV light and ended in 15 min. After polymerization, the cover slip was removed from the surface with forceps and the microcontact BSA imprinted electrode was rinsed with water. The measurements were performed with a capacitive sensor involving the microcontact imprinted electrode inserted in an automated ﬂow injection system developed by Erlandsson et al. [57]. Sensors 2017, 17, 390 14 of 21 Table 3. Capacitive biosensors developed for different targets using microcontact imprinting method. Target Biosensing Method Monomers Dynamic Range LOD Selectivity Stability Ref. Capacitive biosensor Methacrylic acid (MAA); Poly Bovine Serum For human serum albumin >70 assays during 20 8 19 with current ethyleneglycol-dimethacrylate 1.0 10 M–1.0 10 M 1.0 10 M [56] Albumin (BSA) (HSA): 5%; For IgG: 3% 2 months pulse method (PEGDMA) Capacitive biosensor Prostate speciﬁc Selectivity coefﬁcient (k) = 2.27 About same level 17 10 17 with current MAA; EGDMA 2.0 10 M–2.0 10 M 16 10 M [58] antigen (PSA) for HSA, k = 2.02 for IgG during 50 injections pulse method HEMA; (2-Hydroxyethyl Capacitive biosensor K = 3.14 for B. subtilis, k = 3.32 methacrylate), About same level 2 7 1 1 E. coli with current 1.0 10 –1.0 10 CFUmL 70 CFUmL for S. aureus, k = 2.98 for [59] N-methacryloyl-L-histidine during 70 injections pulse method S. paratyphi methyl ester (MAH), EGDMA N-isopropylacrylamide (NIPAm), K = 733.1 for chymotrypsin Capacitive biosensor The loss in performance N,N-methylenebisacryl, amide (chy), k = 10.56 for BSA, 13 7 13 Trypsin with current 1.0 10 M–1.0 10 M 3.0 10 M was about 2% after [60] (MBAAm), Acrylamide, k = 6.50 for lysozyme (Lyz), pulse method 80 analyses Hydroxymethylacrylamide k = 3.46 for cytochrome c (cyt c) Sensors 2017, 17, 390 15 of 21 Sensors 2017, 17, x FOR PEER REVIEW 15 of 21 Capacitive biosensors based on a potential pulse have been used in many applications [1,11,61,62] Capacitive biosensors based on a potential pulse have been used in many applications including microcontact imprinting as seen in Table 3. In this type of capacitive biosensors, a small [1,11,61,62] including microcontact imprinting as seen in Table 3. In this type of capacitive biosensors, potential pulse is applied to the working electrode and the capacitance is measured. This potentiometric a small potential pulse is applied to the working electrode and the capacitance is measured. This pulse concept is sensitive to external electronic disturbances. Therefore, the system is prone to potentiometric pulse concept is sensitive to external electronic disturbances. Therefore, the system is inaccurate measurements and poor baseline stability. The sharp potential pulse may induce damage prone to inaccurate measurements and poor baseline stability. The sharp potential pulse may induce to the surface of the working electrode after a while which results in decreasing response of the damage to the surface of the working electrode after a while which results in decreasing response of electrode and eventually to replacement of the working electrode with a new electrode [57]. Another the electrode and eventually to replacement of the working electrode with a new electrode [57]. alternative way is to use a current pulse method to measure the capacitance at the electrode/solution Another alternative way is to use a current pulse method to measure the capacitance at the interface. Erlandsson et al. [57] developed a new concept to measure capacitance based on a constant electrode/solution interface. Erlandsson et al. [57] developed a new concept to measure capacitance current pulse to the biosensor transducer. In the basis of the principle, the system could be described based on a constant current pulse to the biosensor transducer. In the basis of the principle, the system as a simple resistor-capacitor (RC) circuit model. The schematic representation of the capacitance could be described as a simple resistor-capacitor (RC) circuit model. The schematic representation of measurement via current pulse method is shown in Figure 6. The system consists of: the capacitance measurement via current pulse method is shown in Figure 6. The system consists of: (1) A current source; (1) A current source; (2) An electro-chemical flow-cell which includes three electrodes: the working electrode which is a (2) An electro-chemical ﬂow-cell which includes three electrodes: the working electrode which is thin gold film coated with an insulating layer which functions as a bio-recognition layer to a thin gold film coated with an insulating layer which functions as a bio-recognition layer to immobilize the ligand, the auxiliary and reference electrodes which are made from a platinum immobilize the ligand, the auxiliary and reference electrodes which are made from a platinum wire; wire; (3) A potential differential ampliﬁer; and (3) A potential differential amplifier; and (4) A processor which converts the analogue potential to digital signal. (4) A processor which converts the analogue potential to digital signal. Figure 6. (a) Schematic representation of the capacitive system with current pulse method. The system Figure 6. (a) Schematic representation of the capacitive system with current pulse method. The system is comprised of: (1) current source; (2) flow cell which is connected to the working, reference and is comprised of: (1) current source; (2) ﬂow cell which is connected to the working, reference and auxiliary electrodes; (3) potential differential amplifier; and (4) a processor and ADC where the auxiliary electrodes; (3) potential differential ampliﬁer; and (4) a processor and ADC where the analogue potential is converted to digital signal; (b) A schematic view of Howland current pump used analogue potential is converted to digital signal; (b) A schematic view of Howland current pump used for supplying constant current; (c) Constant current supply to the sensor during the determined time for supplying constant current; (c) Constant current supply to the sensor during the determined time periods to measure the resistance and capacitance; (d) Capacitance is measured every minute and periods to measure the resistance and capacitance; (d) Capacitance is measured every minute and each minute (pulse) contains five sub pulse measurements with 20 ms intervals. (Reproduced from each minute (pulse) contains ﬁve sub pulse measurements with 20 ms intervals. (Reproduced from reference [57] with permission). reference [57] with permission). Sensors 2017, 17, 390 16 of 21 Sensors 2017, 17, 390 16 of 21 When When usin using g current pulse current pulse method, the method, the cap capacitive acitive measurement measurements s were wer perfo e performed rmed with with an an autautomated omated flow-inject ﬂow-injection ion syst system em as sho as shown wn in F in igFigur ure 7 [6 e 73] [63 . ]. Figure 7. Schematic representation of automated flow injection capacitive system. The components Figure 7. Schematic representation of automated ﬂow injection capacitive system. The components shown in the figure are integrated into a box to make a single, portable unit. (Reproduced from shown in the ﬁgure are integrated into a box to make a single, portable unit. (Reproduced from Reference [63] with permission). Reference [63] with permission). In their next study, Ertürk et al. [58] used microcontact imprinting method for detection of an In their next study, Ertürk et al. [58] used microcontact imprinting method for detection important biomarker, prostate specific antigen (PSA), for early detection of prostate cancer with of an important biomarker, prostate speciﬁc antigen (PSA), for early detection of prostate −17 capacitive biosensors. The standard solutions of different concentrations of PSA (2.0 × 10 M–2.0 × cancer with capacitive biosensors. The standard solutions of different concentrations of PSA −10 −15 10 M) were analysed by the system and the LOD value was calculated as 16 × 10 M. HSA and IgG 17 10 (2.0 10 M–2.0 10 M) were analysed by the system and the LOD value was calculated as were used as the competing proteins in order to test the selectivity of the system for PSA. These 16 10 M. HSA and IgG were used as the competing proteins in order to test the selectivity of −1 proteins are normally found in the levels of mg·mL where PSA is found at approximately the system for PSA. These proteins are normally found in the levels of mgmL where PSA is found −1 4.0 ng·mL in human serum. Therefore, when the developed system is tested against proteins at at approximately 4.0 ngmL in human serum. Therefore, when the developed system is tested concentrations of 1 mg/mL which are normally found in one-million-fold higher concentrations, the against proteins at concentrations of 1 mg/mL which are normally found in one-million-fold higher system showed around two times more selectivity for PSA compared to HSA and IgG. In the next concentrations, the system showed around two times more selectivity for PSA compared to HSA and step, the sensitivity and the selectivity of the MIP system were compared with the performance of the IgG. In the next step, the sensitivity and the selectivity of the MIP system were compared with the −14 system based on immobilized Anti-PSA antibody. The LOD was determined to be 12 × 10 M with performance of the system based on immobilized Anti-PSA antibody. The LOD was determined to be the Anti-PSA system. The results showed that the MIP capacitive system was very promising to 12 10 M with the Anti-PSA system. The results showed that the MIP capacitive system was very detect biomarkers which are important for diagnosis of various diseases. The authors compared the promising to detect biomarkers which are important for diagnosis of various diseases. The authors sensitivity of capacitive system for PSA detection with the microcontact imprinted surface plasmon compared the sensitivity of capacitive system for PSA detection with the microcontact imprinted −1 −14 resonance (SPR) system [64]. The LOD value was calculated around 91 pg·mL (18 × 10 M) with surface plasmon resonance (SPR) system [64]. The LOD value was calculated around 91 pgmL the SPR biosensors. This result proves that the capacitive system is approximately 1000 times more (18 10 M) with the SPR biosensors. This result proves that the capacitive system is approximately sensitive than the SPR system. 1000 times more sensitive than the SPR system. Microcontact imprinting method was used for E. coli detection by Idil et al. [59]. The authors Microcontact imprinting method was used for E. coli detection by Idil et al. [59]. The authors used a histidine containing specific monomer (N-methacryloyl-amido-histidine, MAH) as a metal used a histidine containing speciﬁc monomer (N-methacryloyl-amido-histidine, MAH) as a metal 2+ chelating ligand. By using a meal-chelate between MAH and copper(II) (MAH-Cu ), an enhanced 2+ chelating ligand. By using a meal-chelate between MAH and copper(II) (MAH-Cu ), an enhanced selectivity against certain amino acid residues present on the cell wall of E. coli was achieved. selectivity against certain amino acid residues present on the cell wall of E. coli was achieved. N-(hydroxyethyl)methacrylate (HEMA) was used as a functional monomer to make a complex with N-(hydroxyethyl)methacrylate (HEMA) was used as a functional monomer to make a complex 2+ 2+ MAH-Cu (pHEMA-MAH-Cu ) and EGDMA was used as cross-linker. The dynamic range for E. 2+ 2+ with MAH-Cu (pHEMA-MAH-Cu ) and EGDMA was used as cross-linker. The dynamic 2 7 −1 −1 coli detection was between 1.0 × 10 and 1.0 × 10 CFU·mL with a LOD value of 70 CFU·mL . Bacillus 2 7 1 range for E. coli detection was between 1.0 10 and 1.0 10 CFUmL with a LOD value of subtilis, Staphylococcus aureus and Salmonella paratyphi strains were used in selectivity experiments as 70 CFUmL . Bacillus subtilis, Staphylococcus aureus and Salmonella paratyphi strains were used in competing strains. The cross-reactivity ratios were between 24% and 58% against pre-mixed selectivity experiments as competing strains. The cross-reactivity ratios were between 24% and 58% suspensions of all bacterial strains. Even though the system showed cross-reactivity against against pre-mixed suspensions of all bacterial strains. Even though the system showed cross-reactivity competing strains, the ratio was negligible compared to the response of the system against E. coli. against competing strains, the ratio was negligible compared to the response of the system against River water and apple juice were used to show the detection performance of the system from E. coli. River water and apple juice were used to show the detection performance of the system from complex, real samples. Recovery value was found between 81% and 97% for E. coli detection from E. complex, real samples. Recovery value was found between 81% and 97% for E. coli detection from 2 4 −1 coli spiked (1.0 × 10 to 1.0 × 10 CFU·mL ) river water samples. The sensor had potential to monitor E. coli in contaminated water or food supplies. Sensors 2017, 17, 390 17 of 21 2 4 1 E. coli spiked (1.0 10 to 1.0 10 CFUmL ) river water samples. The sensor had potential to monitor E. coli in contaminated water or food supplies. Ertürk et al. [60] used microcontact imprinting method to develop capacitive biosensor for trypsin detection. A monomer solution containing hydroxymethylacrylamide, N-isopropylacrylamide (NIPAm), acrylamide and N,N-methylenebisacrylamide was prepared 0 0 and N,N,N ,N -tetramethylethylenediamine (TEMED, 5%, v/v) and ammonium persulfate (APS, 10%, v/v) were added into this solution. When the modiﬁed gold electrode was treated with monomer solution and Sensors br 2017 ought , 17, 390 into contact with the protein stamp which was carrying immobilized 17 of 21 trypsin on top, the polymerization continued for 3–5 h at room temperature. Schematic representation of Ertürk et al. [60] used microcontact imprinting method to develop capacitive biosensor for microcontact imprinting of trypsin on capacitive electrodes is shown in Figure 8. The dynamic range trypsin detection. A monomer solution containing hydroxymethylacrylamide, 13 7 13 for trypsin detection was between 1.0 10 M and 1.0 10 M with a LOD value of 3.0 10 M. N-isopropylacrylamide (NIPAm), acrylamide and N,N-methylenebisacrylamide was prepared and N,N,N′,N′-tetramethylethylenediamine (TEMED, 5%, v/v) and ammonium persulfate (APS, 10%, v/v) In order to test the selectivity and cross-reactivity of trypsin imprinted capacitive system for trypsin were added into this solution. When the modified gold electrode was treated with monomer solution (MW 23.3 kDa, isoelectric point (pI): 10.1–10.5), chymotrypsin (chy) (MW: 25.6 kDa, pI 8.3), bovine and brought into contact with the protein stamp which was carrying immobilized trypsin on top, the serum albumin (BSA) (MW: 66.5 kDa, pI 4.7), lysozyme (lyz) (MW 14.3 kDa, pI 11.35) and cytochrome c polymerization continued for 3–5 h at room temperature. Schematic representation of microcontact (cyt c) (MW: 12.3 kDa, pI 10.0–10.5) were selected as competing agents. The concentrations of the imprinting of trypsin on capacitive electrodes is shown in Figure 8. The dynamic range for trypsin −13 −7 −13 detection was between 1.0 × 10 M and 1.0 × 10 M with a LOD value of 3.0 × 10 M. In order to test proteins in the selectivity and cross-reactivity experiments were 1.0 mgmL . If it was tested with the the selectivity and cross-reactivity of trypsin imprinted capacitive system for trypsin (MW 23.3 kDa, 1 1 lower concentrations including 1.0 gmL or 1.0 pgmL , the selectivity results would be better isoelectric point (pI): 10.1–10.5), chymotrypsin (chy) (MW: 25.6 kDa, pI 8.3), bovine serum albumin because it would not be possible to detect interfering proteins in these concentration levels. Very low (BSA) (MW: 66.5 kDa, pI 4.7), lysozyme (lyz) (MW 14.3 kDa, pI 11.35) and cytochrome c (cyt c) (MW: afﬁnity of the system towards chy shows that the system can be used successfully for trypsin detection 12.3 kDa, pI 10.0–10.5) were selected as competing agents. The concentrations of the proteins in the −1 selectivity and cross-reactivity experiments were 1.0 mg·mL . If it was tested with the lower from pancreatic secretions where chy and try are found together. When the re-usability of the system −1 −1 concentrations including 1.0 µg·mL or 1.0 pg·mL , the selectivity results would be better because it was tested by monitoring the change in capacitance (-pFcm ) at the same concentration of trypsin would not be possible to detect interfering proteins in these concentration levels. Very low affinity of (10 M), the loss in the detection performance was around 2% after 80 analyses and there was not the system towards chy shows that the system can be used successfully for trypsin detection from any signiﬁcant difference in the performance after storage for two months at 4 C. In the last step, pancreatic secretions where chy and try are found together. When the re-usability of the system was −2 −7 tested by monitoring the change in capacitance (-pF·cm ) at the same concentration of trypsin (10 trypsin activity measured by capacitive system was compared with the trypsin activity measured by M), the loss in the detection performance was around 2% after 80 analyses and there was not any spectrophotometer at 410 nm. One unit of enzyme was deﬁned as the amount of enzyme catalyzing significant difference in the performance after storage for two months at 4 °C. In the last step, trypsin the conversion of one micromole of substrate, N -benzoyl-DL-arginine 4-nitroanilide hydrochloride activity measured by capacitive system was compared with the trypsin activity measured by (BAPNA), at 25 C, pH: 8.1 per minute. The trypsin activity measured spectrophotometrically was spectrophotometer at 410 nm. One unit of enzyme was defined as the amount of enzyme catalyzing 1 1 the conversion of one micromole of substrate, Nα-benzoyl-DL-arginine 4-nitroanilide hydrochloride 9 mUmL where the value was around 8.0 mUmL measured by capacitive system. The results (BAPNA), at 25 °C, pH: 8.1 per minute. The trypsin activity measured spectrophotometrically was 9 showed that there was correlation between two methods. The advantages of the developed method −1 −1 mU·mL where the value was around 8.0 mU·mL measured by capacitive system. The results including detection of trypsin in 20 min, high selectivity towards interfering proteins, high correlation showed that there was correlation between two methods. The advantages of the developed method with the spectrophotometer show that the system might be used successfully for detection of proteases including detection of trypsin in 20 min, high selectivity towards interfering proteins, high correlation with the spectrophotometer show that the system might be used successfully for detection and as a point of care system for diagnosis of pancreatic diseases. of proteases and as a point of care system for diagnosis of pancreatic diseases. Figure 8. Schematic representation of preparation of trypsin imprinted capacitive electrodes using Figure 8. Schematic representation of preparation of trypsin imprinted capacitive electrodes microcontact imprinting procedure: (A) preparation of glass cover slips (protein stamps); (B) using microcontact imprinting procedure: (A) preparation of glass cover slips (protein stamps); preparation of capacitive gold electrodes; and (C) imprinting of trypsin to the electrode surface via (B) preparation of capacitive gold electrodes; and (C) imprinting of trypsin to the electrode surface via microcontact imprinting method. (Reproduced from Reference [60] with permission). microcontact imprinting method. (Reproduced from Reference [60] with permission). Sensors 2017, 17, 390 18 of 21 7. Concluding Remarks From the above, it is obvious that the combination of capacitive biosensors and MIPs offers great possibilities, both with regard to sensitive and selective assays as well as to stable measurements of extended periods of time. 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Capacitive Biosensors and Molecularly Imprinted Electrodes

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Capacitive Biosensors and Molecularly Imprinted Electrodes

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