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A differential capacitive biosensor using polyethylene glycol to overlay the biolayer
Sensors and Actuators B 44 (1997) 578 – 584
A differential capacitive biosensor using polyethylene glycol to overlay the biolayer Helen Berney a,*, John Alderman a, William Lane a, John K. Collins b a
National Microelectronics Research Centre, Uni6ersity College, Cork, Ireland b Microbiology Department, Uni6ersity College, Cork, Ireland
Abstract The feasibility of immunosensors based on capacitance measurements on semiconductor-immobilized antibody-electrolyte heterostructures is being investigated. Capacitance measurements on biosensors succeed only if the successive biomolecular layers grafted onto the heterostructures are sufficiently electrically insulating and retain their recognizing ability (Jaffrezic-Renault et al., Sensors and Actuators B 15–16 (1993) 458–462; Bergveld, Biosensors and Bioelectronics 6 (1991) 55 – 72; Schasfoort et al., Anal. Chim. Acta 238 (1990) 323–329). However, as with other groups, our system showed no reproducible capacitance decrease on addition of antigen. The biolayer was not sufficiently insulating or did not have a suitable dielectric character. Ions were moving through or around this layer causing ‘shorting’ of the system. In an attempt to drive the system toward exclusion of ions from the biolayer, an overlay of non-conducting polymer, in the form of polyethylene glycol (PEG), was added. Our results indicate that PEG forms an insulating layer on a bare chip surface, causing the capacitance to drop. The response was self-consistent and reproducible. On surfaces where there was protein pre-immobilized or pre-adsorbed, the PEG could not form a continuous, integral layer and a smaller drop in capacitance was observed. There is a difference in the response of single layers of protein, where only antibody is immobilized, and triple protein layers, where antibody was immobilized, antigen added and a secondary antibody bound. These results show the possibility of developing a differential capacitive biosensor. © 1997 Elsevier Science S.A. Keywords: Differential capacitive biosensor; Polyethylene glycol; Transferrin
1. Introduction A typical biosensor construction consists of a biological sensing layer combined with a base sensing device that is able to produce an electrical signal when the biological layer interacts with the substance of interest in the applied sample. In order for signal conversion to occur there must be a readily measurable change in a physical or chemical property of the biolayer when recognition takes place. Such changes have been measured using electrochemical, optical, gravimetric, piezoelectric and colorimetric methods [4]. The diversity of naturally occurring biological recognition molecules means that biosensors can be designed to measure a wide range of species of interest. Enzymes and antibodies are often used to provide the biological sensing but tissue slices, whole cells and multienzyme systems
* Corresponding author. Tel.: +353 21 904181; fax: + 353 21 270271; e-mail: [email protected]. 0925-4005/97/$17.00 © 1997 Elsevier Science S.A. All rights reserved. PII S 0 9 2 5 - 4 0 0 5 ( 9 7 ) 0 0 2 3 9 - 6
have also been used as the specific recognition system. The potential applications of these devices are wide and varied. Medical care stands to benefit most clearly and immediately from biosensors in clinical testing, continuous assessment of patient status and development of replacement organs monitoring transplant acceptance/rejection. At present, biosensors are being used to determine food quality and safety, to detect environmental pollutants and in-line process monitoring for various systems [5,6]. The biosensor under development at the NMRC is an immunosensor to detect human transferrin. Transferrin is the major iron binding monomeric protein which is present in all vertebrate species. It carries ferric iron from the intestine, reticuloendothelial system and liver parenchymal cells to all proliferating cells in the body [7]. The laboratory measurement of transferrin provides valuable information in the epidemiological assessment of both iron and protein nutritional status [8]. Abnormal transferrin concentrations are also linked to cancer and renal disorders [9].
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Dividing detection ranges of biological components of interest into a three-tier listing from low range (B10 mg ml − 1), exemplified by blood coagulation factor detection, through medium range (10 – 250 mg ml − 1), typified by transport proteins and tumour markers, to high ( \ 250 mg ml − 1), then transferrin detection is in the upper range. It would be useful to be able to detect 2–4 mg ml − 1 of transferrin in solution. The immunosensor comprises an active biolayer of antibodies (Ab) to human transferrin covalently bound via glutaraldehyde cross-linking to an aminosilane, gaminopropyl triethoxy silane (g-APTES) or 4-amino butyldimethyl methoxysilane (4ABDMMS), on a silicon –silicon dioxide– silicon nitride surface. This was a modification of the immobilization protocol used in Refs. [1,10,11]. A number of groups have tried to measure antibody–antigen interactions using measured capacitance or impedance but the systems have inherent problems with reproducibility, complex analysis of results and measurement time constraints [12 – 15]. Work performed by our own group showed problems related to biolayer surface geometry and morphology [16]. This paper presents a possible solution to some of these problems and a tool to further analyze what happens at semiconductor – biological material –electrolyte interfaces.
1.1. The concept There is a layer of antibody immobilized at the chip surface and, on addition of antigen, the layer becomes thicker and the measured capacitance is expected to decrease. From the electrical viewpoint, the underlying theory of operation of the capacitive immunosensor is quite simple. A capacitor is built consisting of a doped semiconductor (p + -silicon), insulator (35 nm silicon dioxide and 30 nm silicon nitride), antibody and an electrolyte. When an a.c. voltage is applied between the electrolyte and the semiconductor, the insulator and antibody layers act as the dielectric of a capacitor with a capacitance dependent on the layer thickness. If the antigen specific to the antibody is introduced into the electrolyte and binds to the antibody, the dielectric layer will become thicker and so the capacitance will decrease.
1.2. Equation Capacitance is governed by the following equation: C=
oo0A d
where C is the measured capacitance (F), o0 is the effective permittivity (F m − 1), o is the dielectric con-
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stant, A is the area of the exposed chip surface (m2), and d is the layer thickness at the chip surface (m). Values for oox, onit and o0 are as follows: 3.9, 7 and 8.854× 10 − 12 F m − 1. The expected capacitance for a structure consisting solely of semiconductor and insulator (SiO2 35 nm and Si3N4 30 nm) as above is 5 ×10 − 9 F. If the biological membrane values are included as obiol = 20 [17] and layer thickness 20 nm (as measured by ellipsometry), plus an effective layer thickness (antibody electrolyte interface) of 20 nm, the calculated capacitance according to the equation 1/Ctot = 1/Cox + 1/Cnit + 1/Cbiol is 4.6× 10 − 9 F. The measured capacitance of such a structure is 4× 10 − 9 F.
1.3. The problem The immobilized biolayer appeared not to have suitable dielectric properties to respond to the transferrin antigen addition. An immobilized layer of antibody may be sufficiently dense, due to conformational changes or denaturation, to see a modulated capacitive response on addition of antigen. However, many experimental results show that protein layers may be very porous. In the latter case, the layer shows a specific ionic conductivity which may be related to the state of protein charge [2]. It was thought that ions pass through or around the biolayer, providing a ‘shorting’ path [3]. Proteins in their hydrated, solubilized forms have ions associated with them. The solubility of a protein in an aqueous solvent is determined by the distribution of charged hydrophilic and hydrophobic groups on its surface. The charged groups on the surface will interact with ionic groups in the solution [18]. Methods to improve the insulating properties of the protein layers were investigated.
1.4. PEG o6erlay as a possible solution In an attempt to drive the system toward exclusion of ions from the biolayer, an overlay of non-conducting polymer, in the form of polyethylene glycol (PEG), was added. PEG is a linear or branched neutral polyether available in a range of molecular weights. It is widely used in the biomedical industry as a fractional precipitating agent for the purification of proteins from a variety of sources. In aqueous solutions it acts as a highly mobile molecule with a large exclusion volume. It lowers the dielectric constant of the solution and hence its solvating power. Proteins are stearically excluded from regions of the solvent (buffer solution) occupied by PEG and are concentrated until their solubility is exceeded and precipitation occurs. As the concentration of PEG
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increases, the polymers must either shrink and become more compact or interpenetrate, forming a gel-like network [18]. The effect of an antigen – antibody precipitate on the d.c. and a.c. behaviour of an ion-selective field-effect transistor (ISFET) has been studied [19]. However, in that particular case, the precipitate had been preformed and coated onto the ISFET. Our work attempts to precipitate, by exclusion of ions, the immobilized antibody with antigen bound in solution in a non-static way. If PEG is covalently bound to a surface, it alters the electrical nature of that surface, since charges on the surface become buried beneath a viscous hydrated neutral layer. PEG that is added in solution does not bind covalently but forms a quasi-adsorbed layer or overlayer. This allows for a reversible measurement where the PEG solution can be removed by washing. The flexible PEG chains can readily adapt their conformation to fit the surface topology of a protein [20,21].
2. Experimental
2.1. Reagents The basic phosphate-buffered saline solution (PBS), pH 7.4, which was used for the preparation of all other solutions, contained 0.027 M KCl (Avondale Laboratories), 0.0015 M KH2PO4 (Merck), 0.137 M NaCl (Avondale Laboratories) and 0.008 M Na2HPO4 (Riedel–de Haen). All reagents were analytical grade. A washing buffer, PBS/Tween, was prepared by addition of 1 ml of polyoxyethylene sorbitan monolaurate (Tween 20, Sigma) to 1 l of PBS. The blocking buffer was 5 wt.% bovine serum albumin (BSA, Fraction V, 96% purity, Sigma). A solution of 1 mg ml − 1 BSA was used to compare covalently bound to immobilized protein. The mouse monoclonal antibodies (1D2A4) were supplied as ascites by the Microbiology/Virology department of University College, Cork and purified using a protein G column (ImmunoPure Immobilised Protein G, Pierce) chromatography technique, according to the manufacturer’s recommended protocol. The purified antibody solutions were used at a concentration of 100 mg ml − 1 as determined by a Biorad assay. The polyclonal Goat anti-human transferrin antibodies and the transferrin antigen were supplied by Sigma and used at concentrations of 100 mg ml − 1. The polyethylene glycol solution was a 10% w/v PEG 6000 (Sigma) in PBS.
2.2. Chip details The immunochip consists of an low-pressure chemi-
cal vapour deposition (LPCVD) nitride of thickness 30 nm deposited on a CVD oxide of thickness 35 nm deposited on p + silicon. There is an aluminium back contact of thickness 1 mm. This chip allows for the possibility of further integration in the future.
2.3. Vial preparation Polystyrene serum vials were attached to silicon– silicon dioxide–silicon nitride chips, 8× 8 mm, using a UV curable adhesive. The exposed area of the chip once the vial has been attached is 8× 10 − 6 m2. The chips had been silanized using g-aminopropyltriethoxy silane (g-APTES) or 4-aminobutyldimethyl methoxysilane (4-ABDMMS). The silanized chip surface was glutaraldehyde activated prior to overnight incubation with 100 mg ml − 1 of the protein of interest at 4°C. This yielded a covalently bound protein [12,13]. To passively adsorb the protein onto the surface, the glutaraldehyde activation step was omitted. The protein solution was washed out of the vials using PBS/Tween and a 1 h blocking step using 5% BSA in PBS was performed. The vials were equilibrated in a solution of PBS overnight at room temperature prior to testing.
2.4. Experimental set-up The capacitance measurements were carried out using a HP4248A impedance meter at 1 kHz and a signal voltage of 9 100 mV was applied. A platinum wire reference electrode was used. The stability of this electrode at 0 V over a period of 12 h was monitored to ensure that the response is constant with time at the applied voltage. However, because the underlying silicon is heavily doped p + , and CV sweeps have confirmed this, there is very little sensitivity to the bias voltage. The effective contribution of the capacitance of the immunomembrane to the overall capacitance is very much dependent on the choice of operating frequency. Impedance measurements of the devices have shown that the response is almost purely capacitive (measured phase angle u is − 89°) from 20 Hz to 10 kHz. At frequencies greater than 100 kHz, the device response has a resistive component, as seen by a change in u from −89° to some higher value. At our chosen frequency of 1 kHz, we are measuring at a capacitancedominated region. The equilibrated vial contained 400 ml of PBS solution. An initial 1 h drift measurement, to establish system stability, was performed. A series of control measurements was conducted to take into account the effects of light, heat and stirring movement.
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The platinum electrode was removed. A total of 100 ml of PBS was removed and 100 ml of a 10% w/v PEG solution was added to the vial and mixed to give a homogenous solution. The electrode was replaced and the capacitance was measured as a function of time.
3. Results
3.1. Adsorbed 6ersus co6alently bound BSA Initially, an experiment was carried out to establish what effect the addition of PEG had on the different protein conformations, that is covalently immobilized and passively adsorbed. Bovine serum albumin was the protein chosen. There were two control vials that had no protein on the surfaces (8 and 9), four vials with BSA passively adsorbed (1, 2, 3 and 4) and three vials that had BSA covalently bound to the g-APTES silanized surface (5, 6 and 7). As can be seen from Fig. 1, the effect of PEG addition varies depending on the immobilization procedure and surface configuration of the protein. The figure show the effect of PEG addition on pre-immobilized layers. The control vials have no protein on their surfaces. As the PEG coats, forming an insulating layer, there is a change in capacitance. On the surfaces that have had protein immobilized or adsorbed, there was a much smaller change in the capacitance values. Adsorptively bound BSA is in direct contact with the silanized surface. Covalently bound BSA is cross-linked using glutaraldehyde and a matrix is formed. It appeared to be more difficult for the PEG to form an
Fig. 2. Overview of a series of PEG overlay experiments. Vials 1, 2 and 3 have had 1D2A4 covalently immobilized, at 100 ml of 100 mg ml − 1 solution, transferrin added, at 100 ml of 100 mg ml − 1 solution, and goat anti-human transferrin antibody linked to rabbit anti-goat IgG-HRPO as the secondary antibody, at 100 ml of 100 mg ml − 1 solution. Vials 4 and 5 had the single 1D2A4 layer covalently immobilized, at 100 ml of 100 mg ml − 1 solution. Control vial 6 had no protein associated with it. The chip surfaces had been steam activated to provide a suitable number of OH groups for the silanization agent. 4-ABDMMS was used to silanize.
insulating layer over this protein matrix. The covalently bound BSA yielded the smallest capacitive change.
3.2. Triple protein layer 6ersus single protein layer In an attempt to ensure the maximum difference between the responses of the immunovials, a ‘stack’, or triple layer, of proteins was formed which it was postulated would ‘hold off’, or slow down, the PEG overlay to the greatest extent. Fig. 2 show the effect of PEG addition on pre-immobilized layers. The ‘triple’ protein layer consists of a primary monoclonal antibody (1D2A4) immobilized plus antigen (transferrin) plus secondary antibody (goat anti-human transferrin antibody linked to rabbit anti-goat IgG-HRPO). The ‘single’ layer is composed of immobilized monoclonal antibody 1D2A4. There is a reproducible measurable difference between the single and triple layer protein immunovial responses. Control vial 6, which had nothing immobilized on its surface, had the largest drop in capacitance on addition of PEG.
4. Kinetics of reaction Fig. 1. A series of experiments to compare the effect of PEG on BSA that had been adsorbed (1, 2, 3 and 4), at 100 ml of 100 mg ml − 1 solution, on the g-APTES silanized chip surface and BSA that had been covalently immobilized (5, 6 and 7), at 100 ml of 100 mg ml − 1 solution, with the control vials (8 and 9) that had no associated protein on their surfaces.
Regardless of the physical method employed in an attempt to sense the antibody antigen reaction, the performance of the device will mainly be decided by the chemical factors determining the degree of binding [22].
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One of the most important factors is the basic equilibrium and kinetics of the binding reaction at the surface [23]. This reaction is governed by the density and activity of the immobilized species the liquid transducer interface [23,24]. The kinetic parameters of an antigen – antibody interaction are directly related to antibody structure. The selection of the right antibody is of fundamental importance in sensor construction. It is vital to choose an antibody with a strong binding affinity for its antigen and a low dissociation. For the triple protein layer experiments, performed in solid-phase enzyme-linked immunosorbent assay (ELISA) measurements, the monoclonal antibody 1D2A4 acted as the primary layer. To this layer was bound antigen and polyclonal antibody as a secondary layer. The advantage of using an antibody as the primary layer to trap antigen is that the antigen tends to preserve the immunoactivity of the antigen. Parameters which play a role in solid-phase-based methods are diffusion, stearic hindrance of the immobilized molecules and alteration of the binding site due to conformational changes. Polyclonal antibodies possess an inherent heterogeneity in that the antibodies in a particular sample are not identical. An increase in external mass transfer limitations decreases the saturation level of the antigen close to the surface and also the rate of antigen attachment to the antibody covalently or non-covalently bound to the surface [25,26]. It is generally accepted that antibody – antigen reactions have time constants in the order of hours for equilibrium to be reached. However, if the rate of association and dissociation are monitored, the maximum amount of antibody is bound within 15 min, and if the dissociation constant is low this value remains steady [27]. For our system, the protein layers were pre-adsorbed so no effect of reaction kinetics were monitored.
5. Possible mechanisms of PEG action On the protein-free surface the PEG forms an insulating layer which causes the capacitance to drop. When there is protein at the surface, the PEG may fill in the spaces between the protein molecules and thus form the required insulating layer or it may coat over the proteins, reducing the ion movement through the layer. The protein appeared to slow down or prevent the complete PEG insulating layer forming. Although the initial change was swifter than for the non-protein surface, an equilibrium was reached and the percentage reduction was smaller.
6. Discussion There is a possibility of using this technique to establish a differential immunosensor, where difference in response between a single antibody layer (to which the antigen of interest does not bind) and an antibody layer to which antigen has bound. This has the advantage that the background conditions, which in a direct measurement have to be corrected, are taken into account in the reference sensor. H. Maupas et al. at Eurosensors X [28] showed a system based on a differential sensor based on impedance measurements. In this paper he shows a test time of up to 3 h. Classical ELISA techniques may take up to 24 h to establish a result and involve many steps (immobilization, blocking, binding washing, addition of labelled antibody, washing, addition of substrate and measurement). Ideally, a differential system would be established in a flow-through cell where after addition of sample the PEG is added and the differential response monitored. Within 0.5–1 h, the differential signal should yield the result. However, this presupposes a relationship between amount of antigen added and measured capacitance drop. At the moment, the system is used so that there is a maximum difference between the reference (nothing on the chip surface), the test (antibody plus antigen), and test 2 (antibody plus antigen plus antibody). It is only viable to detect the presence or absence of a particular antigen, and while useful for certain types of analysis, the goal is to develop a quantitative measure of antigen.
7. Conclusions There is a measurable difference between the single and triple layer proteins, which indicates a possibility of using a PEG overlay system in the development of a differential capacitive biosensor. The response seen on addition of PEG to the vials is self-consistent and reproducible. The underlying immobilized antibody confers specificity and the PEG allow a capacitive response to be measured.
8. Future work More quantitative information is required on the capacitance change with respect to the amount of protein immobilized on the surface. A calibration of total protein versus capacitance drop would be a useful measure of the detection limits of the system.
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Acknowledgements This research was part funded by the EU within the frame of the ESPRIT 6240 project ‘Biological/Chemical ASICs for Smart Sensors in Medical and Environmental Monitoring’.
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Biographies Helen Berney received a B.Sc. in Biotechnology from Dublin City University in 1991. She joined the Cancer Research Campaign Laboratory, Nottingham, England in 1991 where she worked producing, characterizing and developing monoclonal antibodies to dust mite for use in treatment of asthma. In 1993, she joined the National Microelectronics Research Centre (NMRC), Cork, Ireland as part of a team developing biological and chemical sensors for environmental and medical monitoring. John Alderman received a B.Sc. in Chemistry from Bristol University in 1978 and an M.Sc. in Microelectronics from Bangor University in 1982. He joined Plessey Research (Caswell), England in 1978 and after initial work on 5 mm metal gate NMOS moved on to polygate CCDs and NMOS. He started work on SOI developments in the early 1980s, working on recrystallized, porous and SIMOX material, latterly concentrating on sub-micron CMOS devices and the comparison to bulk devices. He joined the NMRC, Cork in 1990, initially to continue SOI activities within the ESPRIT
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SUB SOITEC project. He is also involved in using e-beam direct write for ultra-fine geometries and in the Smart sensor/Microsystems area with biosensor and environmental monitoring under investigation in the ESPRIT B-ASICs project. He is also involved in a basic and long-term project BARMINT investigating primarily MST packaging aspects. He has several patents filed and has published and presented over 60 papers in the technical literature either as a sole or joint author in the above areas of interest. William A. Lane received the B.Sc. and M.Sc. degrees from the University of Toronto, Canada in 1977 and 1980, respectively. His master’s work was on the development of power VMOS devices and power ICs. In 1987 he was awarded the Ph.D. degree from the National University of Ireland for work on precision analogue polysilicon resistors in CMOS processed. From 1980 to 1987 he worked at the NMRC, Cork, Ireland on the development of bulk CMOS and BICMOS processes, in modelling of these processes for circuit simulation and in the development of SOI devices and processes. In 1987 he became Process Development Manager for Analog Devices B.V. in Limerick, Ireland and was responsible for the development of several BiCMOS processes for mixed signal application
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and an SOI-based Analog switch process. In 1990 he rejoined the NMRC and his current research interests are in the area of chemical and physical sensors and power devices. He has authored or co-authored more than 50 papers and is a member of the IEEE. John K. Collins received a B.Sc. in Biochemistry/Microbiology from University College, Cork in 1970 and a Ph.D. from the National University of Ireland in 1974. He worked as a post-doctoral research fellow in the departments of Anatomy and Pharmacology in Case Western Research, University Medical School, Cleveland, OH from 1974 to 1975. From 1975 to 1976 he was employed as a post-doctoral research fellow by the department of Microbiology in the State University of New York. From 1976 to 1978 he worked in the departments of Bacteriology and Pathology at the University of California. He was president of the Irish Association for Cancer Research from 1983 to 1993 and a council member of the European Association for Cancer Research from 1985 to 1993. At present he is associate professor of Medicine and Microbiology in University College, Cork. His research interests include inflammatory bowel disease, human tumour biology and immunology, virology and sensors for the medical field. He has in excess of 50 original publications.
A differential capacitive biosensor using polyethylene glycol to overlay the biolayer Helen Berney a,*, John Alderman a, William Lane a, John K. Collins b a
National Microelectronics Research Centre, Uni6ersity College, Cork, Ireland b Microbiology Department, Uni6ersity College, Cork, Ireland
Abstract The feasibility of immunosensors based on capacitance measurements on semiconductor-immobilized antibody-electrolyte heterostructures is being investigated. Capacitance measurements on biosensors succeed only if the successive biomolecular layers grafted onto the heterostructures are sufficiently electrically insulating and retain their recognizing ability (Jaffrezic-Renault et al., Sensors and Actuators B 15–16 (1993) 458–462; Bergveld, Biosensors and Bioelectronics 6 (1991) 55 – 72; Schasfoort et al., Anal. Chim. Acta 238 (1990) 323–329). However, as with other groups, our system showed no reproducible capacitance decrease on addition of antigen. The biolayer was not sufficiently insulating or did not have a suitable dielectric character. Ions were moving through or around this layer causing ‘shorting’ of the system. In an attempt to drive the system toward exclusion of ions from the biolayer, an overlay of non-conducting polymer, in the form of polyethylene glycol (PEG), was added. Our results indicate that PEG forms an insulating layer on a bare chip surface, causing the capacitance to drop. The response was self-consistent and reproducible. On surfaces where there was protein pre-immobilized or pre-adsorbed, the PEG could not form a continuous, integral layer and a smaller drop in capacitance was observed. There is a difference in the response of single layers of protein, where only antibody is immobilized, and triple protein layers, where antibody was immobilized, antigen added and a secondary antibody bound. These results show the possibility of developing a differential capacitive biosensor. © 1997 Elsevier Science S.A. Keywords: Differential capacitive biosensor; Polyethylene glycol; Transferrin
1. Introduction A typical biosensor construction consists of a biological sensing layer combined with a base sensing device that is able to produce an electrical signal when the biological layer interacts with the substance of interest in the applied sample. In order for signal conversion to occur there must be a readily measurable change in a physical or chemical property of the biolayer when recognition takes place. Such changes have been measured using electrochemical, optical, gravimetric, piezoelectric and colorimetric methods [4]. The diversity of naturally occurring biological recognition molecules means that biosensors can be designed to measure a wide range of species of interest. Enzymes and antibodies are often used to provide the biological sensing but tissue slices, whole cells and multienzyme systems
* Corresponding author. Tel.: +353 21 904181; fax: + 353 21 270271; e-mail: [email protected]. 0925-4005/97/$17.00 © 1997 Elsevier Science S.A. All rights reserved. PII S 0 9 2 5 - 4 0 0 5 ( 9 7 ) 0 0 2 3 9 - 6
have also been used as the specific recognition system. The potential applications of these devices are wide and varied. Medical care stands to benefit most clearly and immediately from biosensors in clinical testing, continuous assessment of patient status and development of replacement organs monitoring transplant acceptance/rejection. At present, biosensors are being used to determine food quality and safety, to detect environmental pollutants and in-line process monitoring for various systems [5,6]. The biosensor under development at the NMRC is an immunosensor to detect human transferrin. Transferrin is the major iron binding monomeric protein which is present in all vertebrate species. It carries ferric iron from the intestine, reticuloendothelial system and liver parenchymal cells to all proliferating cells in the body [7]. The laboratory measurement of transferrin provides valuable information in the epidemiological assessment of both iron and protein nutritional status [8]. Abnormal transferrin concentrations are also linked to cancer and renal disorders [9].
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Dividing detection ranges of biological components of interest into a three-tier listing from low range (B10 mg ml − 1), exemplified by blood coagulation factor detection, through medium range (10 – 250 mg ml − 1), typified by transport proteins and tumour markers, to high ( \ 250 mg ml − 1), then transferrin detection is in the upper range. It would be useful to be able to detect 2–4 mg ml − 1 of transferrin in solution. The immunosensor comprises an active biolayer of antibodies (Ab) to human transferrin covalently bound via glutaraldehyde cross-linking to an aminosilane, gaminopropyl triethoxy silane (g-APTES) or 4-amino butyldimethyl methoxysilane (4ABDMMS), on a silicon –silicon dioxide– silicon nitride surface. This was a modification of the immobilization protocol used in Refs. [1,10,11]. A number of groups have tried to measure antibody–antigen interactions using measured capacitance or impedance but the systems have inherent problems with reproducibility, complex analysis of results and measurement time constraints [12 – 15]. Work performed by our own group showed problems related to biolayer surface geometry and morphology [16]. This paper presents a possible solution to some of these problems and a tool to further analyze what happens at semiconductor – biological material –electrolyte interfaces.
1.1. The concept There is a layer of antibody immobilized at the chip surface and, on addition of antigen, the layer becomes thicker and the measured capacitance is expected to decrease. From the electrical viewpoint, the underlying theory of operation of the capacitive immunosensor is quite simple. A capacitor is built consisting of a doped semiconductor (p + -silicon), insulator (35 nm silicon dioxide and 30 nm silicon nitride), antibody and an electrolyte. When an a.c. voltage is applied between the electrolyte and the semiconductor, the insulator and antibody layers act as the dielectric of a capacitor with a capacitance dependent on the layer thickness. If the antigen specific to the antibody is introduced into the electrolyte and binds to the antibody, the dielectric layer will become thicker and so the capacitance will decrease.
1.2. Equation Capacitance is governed by the following equation: C=
oo0A d
where C is the measured capacitance (F), o0 is the effective permittivity (F m − 1), o is the dielectric con-
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stant, A is the area of the exposed chip surface (m2), and d is the layer thickness at the chip surface (m). Values for oox, onit and o0 are as follows: 3.9, 7 and 8.854× 10 − 12 F m − 1. The expected capacitance for a structure consisting solely of semiconductor and insulator (SiO2 35 nm and Si3N4 30 nm) as above is 5 ×10 − 9 F. If the biological membrane values are included as obiol = 20 [17] and layer thickness 20 nm (as measured by ellipsometry), plus an effective layer thickness (antibody electrolyte interface) of 20 nm, the calculated capacitance according to the equation 1/Ctot = 1/Cox + 1/Cnit + 1/Cbiol is 4.6× 10 − 9 F. The measured capacitance of such a structure is 4× 10 − 9 F.
1.3. The problem The immobilized biolayer appeared not to have suitable dielectric properties to respond to the transferrin antigen addition. An immobilized layer of antibody may be sufficiently dense, due to conformational changes or denaturation, to see a modulated capacitive response on addition of antigen. However, many experimental results show that protein layers may be very porous. In the latter case, the layer shows a specific ionic conductivity which may be related to the state of protein charge [2]. It was thought that ions pass through or around the biolayer, providing a ‘shorting’ path [3]. Proteins in their hydrated, solubilized forms have ions associated with them. The solubility of a protein in an aqueous solvent is determined by the distribution of charged hydrophilic and hydrophobic groups on its surface. The charged groups on the surface will interact with ionic groups in the solution [18]. Methods to improve the insulating properties of the protein layers were investigated.
1.4. PEG o6erlay as a possible solution In an attempt to drive the system toward exclusion of ions from the biolayer, an overlay of non-conducting polymer, in the form of polyethylene glycol (PEG), was added. PEG is a linear or branched neutral polyether available in a range of molecular weights. It is widely used in the biomedical industry as a fractional precipitating agent for the purification of proteins from a variety of sources. In aqueous solutions it acts as a highly mobile molecule with a large exclusion volume. It lowers the dielectric constant of the solution and hence its solvating power. Proteins are stearically excluded from regions of the solvent (buffer solution) occupied by PEG and are concentrated until their solubility is exceeded and precipitation occurs. As the concentration of PEG
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increases, the polymers must either shrink and become more compact or interpenetrate, forming a gel-like network [18]. The effect of an antigen – antibody precipitate on the d.c. and a.c. behaviour of an ion-selective field-effect transistor (ISFET) has been studied [19]. However, in that particular case, the precipitate had been preformed and coated onto the ISFET. Our work attempts to precipitate, by exclusion of ions, the immobilized antibody with antigen bound in solution in a non-static way. If PEG is covalently bound to a surface, it alters the electrical nature of that surface, since charges on the surface become buried beneath a viscous hydrated neutral layer. PEG that is added in solution does not bind covalently but forms a quasi-adsorbed layer or overlayer. This allows for a reversible measurement where the PEG solution can be removed by washing. The flexible PEG chains can readily adapt their conformation to fit the surface topology of a protein [20,21].
2. Experimental
2.1. Reagents The basic phosphate-buffered saline solution (PBS), pH 7.4, which was used for the preparation of all other solutions, contained 0.027 M KCl (Avondale Laboratories), 0.0015 M KH2PO4 (Merck), 0.137 M NaCl (Avondale Laboratories) and 0.008 M Na2HPO4 (Riedel–de Haen). All reagents were analytical grade. A washing buffer, PBS/Tween, was prepared by addition of 1 ml of polyoxyethylene sorbitan monolaurate (Tween 20, Sigma) to 1 l of PBS. The blocking buffer was 5 wt.% bovine serum albumin (BSA, Fraction V, 96% purity, Sigma). A solution of 1 mg ml − 1 BSA was used to compare covalently bound to immobilized protein. The mouse monoclonal antibodies (1D2A4) were supplied as ascites by the Microbiology/Virology department of University College, Cork and purified using a protein G column (ImmunoPure Immobilised Protein G, Pierce) chromatography technique, according to the manufacturer’s recommended protocol. The purified antibody solutions were used at a concentration of 100 mg ml − 1 as determined by a Biorad assay. The polyclonal Goat anti-human transferrin antibodies and the transferrin antigen were supplied by Sigma and used at concentrations of 100 mg ml − 1. The polyethylene glycol solution was a 10% w/v PEG 6000 (Sigma) in PBS.
2.2. Chip details The immunochip consists of an low-pressure chemi-
cal vapour deposition (LPCVD) nitride of thickness 30 nm deposited on a CVD oxide of thickness 35 nm deposited on p + silicon. There is an aluminium back contact of thickness 1 mm. This chip allows for the possibility of further integration in the future.
2.3. Vial preparation Polystyrene serum vials were attached to silicon– silicon dioxide–silicon nitride chips, 8× 8 mm, using a UV curable adhesive. The exposed area of the chip once the vial has been attached is 8× 10 − 6 m2. The chips had been silanized using g-aminopropyltriethoxy silane (g-APTES) or 4-aminobutyldimethyl methoxysilane (4-ABDMMS). The silanized chip surface was glutaraldehyde activated prior to overnight incubation with 100 mg ml − 1 of the protein of interest at 4°C. This yielded a covalently bound protein [12,13]. To passively adsorb the protein onto the surface, the glutaraldehyde activation step was omitted. The protein solution was washed out of the vials using PBS/Tween and a 1 h blocking step using 5% BSA in PBS was performed. The vials were equilibrated in a solution of PBS overnight at room temperature prior to testing.
2.4. Experimental set-up The capacitance measurements were carried out using a HP4248A impedance meter at 1 kHz and a signal voltage of 9 100 mV was applied. A platinum wire reference electrode was used. The stability of this electrode at 0 V over a period of 12 h was monitored to ensure that the response is constant with time at the applied voltage. However, because the underlying silicon is heavily doped p + , and CV sweeps have confirmed this, there is very little sensitivity to the bias voltage. The effective contribution of the capacitance of the immunomembrane to the overall capacitance is very much dependent on the choice of operating frequency. Impedance measurements of the devices have shown that the response is almost purely capacitive (measured phase angle u is − 89°) from 20 Hz to 10 kHz. At frequencies greater than 100 kHz, the device response has a resistive component, as seen by a change in u from −89° to some higher value. At our chosen frequency of 1 kHz, we are measuring at a capacitancedominated region. The equilibrated vial contained 400 ml of PBS solution. An initial 1 h drift measurement, to establish system stability, was performed. A series of control measurements was conducted to take into account the effects of light, heat and stirring movement.
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The platinum electrode was removed. A total of 100 ml of PBS was removed and 100 ml of a 10% w/v PEG solution was added to the vial and mixed to give a homogenous solution. The electrode was replaced and the capacitance was measured as a function of time.
3. Results
3.1. Adsorbed 6ersus co6alently bound BSA Initially, an experiment was carried out to establish what effect the addition of PEG had on the different protein conformations, that is covalently immobilized and passively adsorbed. Bovine serum albumin was the protein chosen. There were two control vials that had no protein on the surfaces (8 and 9), four vials with BSA passively adsorbed (1, 2, 3 and 4) and three vials that had BSA covalently bound to the g-APTES silanized surface (5, 6 and 7). As can be seen from Fig. 1, the effect of PEG addition varies depending on the immobilization procedure and surface configuration of the protein. The figure show the effect of PEG addition on pre-immobilized layers. The control vials have no protein on their surfaces. As the PEG coats, forming an insulating layer, there is a change in capacitance. On the surfaces that have had protein immobilized or adsorbed, there was a much smaller change in the capacitance values. Adsorptively bound BSA is in direct contact with the silanized surface. Covalently bound BSA is cross-linked using glutaraldehyde and a matrix is formed. It appeared to be more difficult for the PEG to form an
Fig. 2. Overview of a series of PEG overlay experiments. Vials 1, 2 and 3 have had 1D2A4 covalently immobilized, at 100 ml of 100 mg ml − 1 solution, transferrin added, at 100 ml of 100 mg ml − 1 solution, and goat anti-human transferrin antibody linked to rabbit anti-goat IgG-HRPO as the secondary antibody, at 100 ml of 100 mg ml − 1 solution. Vials 4 and 5 had the single 1D2A4 layer covalently immobilized, at 100 ml of 100 mg ml − 1 solution. Control vial 6 had no protein associated with it. The chip surfaces had been steam activated to provide a suitable number of OH groups for the silanization agent. 4-ABDMMS was used to silanize.
insulating layer over this protein matrix. The covalently bound BSA yielded the smallest capacitive change.
3.2. Triple protein layer 6ersus single protein layer In an attempt to ensure the maximum difference between the responses of the immunovials, a ‘stack’, or triple layer, of proteins was formed which it was postulated would ‘hold off’, or slow down, the PEG overlay to the greatest extent. Fig. 2 show the effect of PEG addition on pre-immobilized layers. The ‘triple’ protein layer consists of a primary monoclonal antibody (1D2A4) immobilized plus antigen (transferrin) plus secondary antibody (goat anti-human transferrin antibody linked to rabbit anti-goat IgG-HRPO). The ‘single’ layer is composed of immobilized monoclonal antibody 1D2A4. There is a reproducible measurable difference between the single and triple layer protein immunovial responses. Control vial 6, which had nothing immobilized on its surface, had the largest drop in capacitance on addition of PEG.
4. Kinetics of reaction Fig. 1. A series of experiments to compare the effect of PEG on BSA that had been adsorbed (1, 2, 3 and 4), at 100 ml of 100 mg ml − 1 solution, on the g-APTES silanized chip surface and BSA that had been covalently immobilized (5, 6 and 7), at 100 ml of 100 mg ml − 1 solution, with the control vials (8 and 9) that had no associated protein on their surfaces.
Regardless of the physical method employed in an attempt to sense the antibody antigen reaction, the performance of the device will mainly be decided by the chemical factors determining the degree of binding [22].
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One of the most important factors is the basic equilibrium and kinetics of the binding reaction at the surface [23]. This reaction is governed by the density and activity of the immobilized species the liquid transducer interface [23,24]. The kinetic parameters of an antigen – antibody interaction are directly related to antibody structure. The selection of the right antibody is of fundamental importance in sensor construction. It is vital to choose an antibody with a strong binding affinity for its antigen and a low dissociation. For the triple protein layer experiments, performed in solid-phase enzyme-linked immunosorbent assay (ELISA) measurements, the monoclonal antibody 1D2A4 acted as the primary layer. To this layer was bound antigen and polyclonal antibody as a secondary layer. The advantage of using an antibody as the primary layer to trap antigen is that the antigen tends to preserve the immunoactivity of the antigen. Parameters which play a role in solid-phase-based methods are diffusion, stearic hindrance of the immobilized molecules and alteration of the binding site due to conformational changes. Polyclonal antibodies possess an inherent heterogeneity in that the antibodies in a particular sample are not identical. An increase in external mass transfer limitations decreases the saturation level of the antigen close to the surface and also the rate of antigen attachment to the antibody covalently or non-covalently bound to the surface [25,26]. It is generally accepted that antibody – antigen reactions have time constants in the order of hours for equilibrium to be reached. However, if the rate of association and dissociation are monitored, the maximum amount of antibody is bound within 15 min, and if the dissociation constant is low this value remains steady [27]. For our system, the protein layers were pre-adsorbed so no effect of reaction kinetics were monitored.
5. Possible mechanisms of PEG action On the protein-free surface the PEG forms an insulating layer which causes the capacitance to drop. When there is protein at the surface, the PEG may fill in the spaces between the protein molecules and thus form the required insulating layer or it may coat over the proteins, reducing the ion movement through the layer. The protein appeared to slow down or prevent the complete PEG insulating layer forming. Although the initial change was swifter than for the non-protein surface, an equilibrium was reached and the percentage reduction was smaller.
6. Discussion There is a possibility of using this technique to establish a differential immunosensor, where difference in response between a single antibody layer (to which the antigen of interest does not bind) and an antibody layer to which antigen has bound. This has the advantage that the background conditions, which in a direct measurement have to be corrected, are taken into account in the reference sensor. H. Maupas et al. at Eurosensors X [28] showed a system based on a differential sensor based on impedance measurements. In this paper he shows a test time of up to 3 h. Classical ELISA techniques may take up to 24 h to establish a result and involve many steps (immobilization, blocking, binding washing, addition of labelled antibody, washing, addition of substrate and measurement). Ideally, a differential system would be established in a flow-through cell where after addition of sample the PEG is added and the differential response monitored. Within 0.5–1 h, the differential signal should yield the result. However, this presupposes a relationship between amount of antigen added and measured capacitance drop. At the moment, the system is used so that there is a maximum difference between the reference (nothing on the chip surface), the test (antibody plus antigen), and test 2 (antibody plus antigen plus antibody). It is only viable to detect the presence or absence of a particular antigen, and while useful for certain types of analysis, the goal is to develop a quantitative measure of antigen.
7. Conclusions There is a measurable difference between the single and triple layer proteins, which indicates a possibility of using a PEG overlay system in the development of a differential capacitive biosensor. The response seen on addition of PEG to the vials is self-consistent and reproducible. The underlying immobilized antibody confers specificity and the PEG allow a capacitive response to be measured.
8. Future work More quantitative information is required on the capacitance change with respect to the amount of protein immobilized on the surface. A calibration of total protein versus capacitance drop would be a useful measure of the detection limits of the system.
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Acknowledgements This research was part funded by the EU within the frame of the ESPRIT 6240 project ‘Biological/Chemical ASICs for Smart Sensors in Medical and Environmental Monitoring’.
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Biographies Helen Berney received a B.Sc. in Biotechnology from Dublin City University in 1991. She joined the Cancer Research Campaign Laboratory, Nottingham, England in 1991 where she worked producing, characterizing and developing monoclonal antibodies to dust mite for use in treatment of asthma. In 1993, she joined the National Microelectronics Research Centre (NMRC), Cork, Ireland as part of a team developing biological and chemical sensors for environmental and medical monitoring. John Alderman received a B.Sc. in Chemistry from Bristol University in 1978 and an M.Sc. in Microelectronics from Bangor University in 1982. He joined Plessey Research (Caswell), England in 1978 and after initial work on 5 mm metal gate NMOS moved on to polygate CCDs and NMOS. He started work on SOI developments in the early 1980s, working on recrystallized, porous and SIMOX material, latterly concentrating on sub-micron CMOS devices and the comparison to bulk devices. He joined the NMRC, Cork in 1990, initially to continue SOI activities within the ESPRIT
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SUB SOITEC project. He is also involved in using e-beam direct write for ultra-fine geometries and in the Smart sensor/Microsystems area with biosensor and environmental monitoring under investigation in the ESPRIT B-ASICs project. He is also involved in a basic and long-term project BARMINT investigating primarily MST packaging aspects. He has several patents filed and has published and presented over 60 papers in the technical literature either as a sole or joint author in the above areas of interest. William A. Lane received the B.Sc. and M.Sc. degrees from the University of Toronto, Canada in 1977 and 1980, respectively. His master’s work was on the development of power VMOS devices and power ICs. In 1987 he was awarded the Ph.D. degree from the National University of Ireland for work on precision analogue polysilicon resistors in CMOS processed. From 1980 to 1987 he worked at the NMRC, Cork, Ireland on the development of bulk CMOS and BICMOS processes, in modelling of these processes for circuit simulation and in the development of SOI devices and processes. In 1987 he became Process Development Manager for Analog Devices B.V. in Limerick, Ireland and was responsible for the development of several BiCMOS processes for mixed signal application
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and an SOI-based Analog switch process. In 1990 he rejoined the NMRC and his current research interests are in the area of chemical and physical sensors and power devices. He has authored or co-authored more than 50 papers and is a member of the IEEE. John K. Collins received a B.Sc. in Biochemistry/Microbiology from University College, Cork in 1970 and a Ph.D. from the National University of Ireland in 1974. He worked as a post-doctoral research fellow in the departments of Anatomy and Pharmacology in Case Western Research, University Medical School, Cleveland, OH from 1974 to 1975. From 1975 to 1976 he was employed as a post-doctoral research fellow by the department of Microbiology in the State University of New York. From 1976 to 1978 he worked in the departments of Bacteriology and Pathology at the University of California. He was president of the Irish Association for Cancer Research from 1983 to 1993 and a council member of the European Association for Cancer Research from 1985 to 1993. At present he is associate professor of Medicine and Microbiology in University College, Cork. His research interests include inflammatory bowel disease, human tumour biology and immunology, virology and sensors for the medical field. He has in excess of 50 original publications.