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Fabrication of BioFET linear array for detection of protein interactions
Microelectronic Engineering 87 (2010) 753–755
Contents lists available at ScienceDirect
Microelectronic Engineering journal homepage: www.elsevier.com/locate/mee
Fabrication of BioFET linear array for detection of protein interactions Ling Wang a,*, Pedro Estrela b,1, Ejaz Huq a, Peng Li b, Stephen Thomas a, Paul Ko Ferrigno c, Debjani Paul b, Paul Adkin a, Piero Migliorato b a
Technology Department, Rutherford Appleton Laboratory, STFC, Oxfordshire, OX11 0QX, UK Department of Engineering, University of Cambridge, Cambridge CB3 0FA, UK c Leeds Institute of Molecular Medicine, St. James’ University Hospital, Leeds LS9 7TF, UK b
a r t i c l e
i n f o
Article history: Received 14 September 2009 Received in revised form 26 November 2009 Accepted 26 November 2009 Available online 2 December 2009 Keywords: BioFETs Protein interaction SU-8
a b s t r a c t An extended-gate MOSFETs (metal–oxide-semiconductor field-effect transistors) based biosensing linear array has been fabricated for label-free protein interaction detection. The device was realized using a combination of very low leakage current MOSFET transistors and an external gate where the chemical reaction would take place. Peptide aptamers that recognize cyclin-dependent kinase (CDK), a protein cancer marker, were used as a biological test system. The test results showed a high sensitive in the detection of CDK. Ó 2009 Elsevier B.V. All rights reserved.
1. Introduction There has been a growing interest in silicon-based field-effect biosensors since Bergveld first investigated the use of ion-sensitivity field-effect transistors (ISFETs) three decades ago [1,2]. Biologically active field-effect transistors (BioFETs) are miniaturized silicon-based electrochemical biosensors that can detect biomolecular interactions occurring at the gates of MOSFETs. Compared to the conventional optical and mass spectrometry detection methods, BioFETs provide fast, label-free, high sensitivity and direct electronic signal readout [3,4]. FET-based biosensors have the potential for developing high-throughput label-free microarrays for applications in clinical diagnostics, drug discovery, food monitoring, and bioterrorist security inspection. In this paper we present a convenient approach for the fabrication of a linear array extended-gate BioFETs. MOSFETs were used as electrical transducers with the gates connected to external gold sensing pads where the electrolyte is introduced and biomolecular interactions take place. This approach allows an easy separation between the sensing area and the transistor in order to reduce any risk of leakage. It also enables an easy way to define and modify the sensing area and shape, and avoiding post-processing on MOSFET itself. The operation of the BioFET detection is based on the immobilization of probe molecules on the extended gate, * Corresponding author. Tel.: +44 1235 445829; fax: +44 1235 446283. E-mail address: [email protected] (L. Wang). 1 Present address: Department of Electronic and Electrical Engineering, University of Bath, Bath BA2 7AY, UK. 0167-9317/$ - see front matter Ó 2009 Elsevier B.V. All rights reserved. doi:10.1016/j.mee.2009.11.148
which is placed in contact with the electrolyte. The gate voltage is applied via a reference electrode immersed in the electrolyte (see Fig. 1). Upon a biomolecular interaction, the change in charge distribution of the bio-layer induces a variation in the potential drop at the interface with the electrolyte, which in turn modulates the gate voltage. Peptide aptamers are artificial recognition molecules made from an insertion of a variable peptide sequence into a constant scaffold protein [5], which allow specific protein targets to bind and get recognized. With the advantages of excellent recognition specificity and high binding affinity compared to antibodies, the peptide aptamers have many potential applications in protein detections. In this paper, the interaction of different peptide aptamers with CDK2 proteins was detected by measuring the transfer characteristics of the MOSFETs.
2. Fabrication The MOSFETs were fabricated at IMEC (Belgium), with a number of n-type and p-type transistors on each chip. The gate lengths of the MOSFETs ranged from 0.7 lm to 2.0 lm, and widths ranged from 3860 lm to 1340 lm with a constant gate area. The MOSFETs were fabricated with 17 nm thick of gate oxide. The operation voltage is 3.3–5 V with a saturation current of 358 lA/lm. There are no electrostatic discharge protection elements in the MOSFET devices in order to minimize the leakage current. The leakage current is a critical concern in this type of detection, since a very small charge change during protein interaction will be detected. A calculation
754
L. Wang et al. / Microelectronic Engineering 87 (2010) 753–755
Reference Electrode
VGS
VDS
Extended gate
n+
n+
Source
Drain P-type Si
Fig. 1. Schematic structure of BioFET detection of protein.
result showed that the gate leakage current must be lower than 10 14 A. The linear array sensing pads were designed 500 500 lm gold pads separated by 1 mm for ease of fabrication. The gold sensing pads were fabricated on glass substrates using a lift-off process. A double layer of photoresist process was used for high quality gold pad edge profile, especially for the very long and thin gold lines, which connect the sensing pads and the wire-bond pads. A layer of LOR (lift-off resist) 20B was spin coated and baked at 180 °C on a hot plate for 10 min. followed by a layer of JSR photoresist spin coated and baked, which resulting in a 2 lm thick LOR and 1.2 lm thick JSR resist. After exposure on a Karl Suss Mask Aligner [UV400 light] the JSR photoresist was developed and hard baked. The LOR was then etched in dilute AZ400 K for about 2 min. to obtain an undercut profile. An SEM picture of the cross-section is shown in Fig. 2. The gold layer (100 nm thick) was deposited in an electron beam evaporator (SV2000) with a thin layer of chrome (10 nm thick) as seed layer. In order to obtain a very smooth gold surface which is required for the protein binding, the gold film was deposited with a low deposition rate of 0.5 Å/s. The lift-off process was achieved in acetone followed by soaking in dilute AZ400 K to remove the LOR. SU-8 epoxy-based photoresist was used as passivation layer to cover the whole surface except the sensing pads and wire-bond
pads. The advantages of the SU-8 include its better biocompatibility compared to silicon dioxide, easier to pattern and good insulation properties. The disadvantage of SU-8 is its weak adhesion to glass surface and stress problem. An investigation for the SU-8 process was performed to optimize the thickness, exposure dose and baking conditions. A series of tests for optimization of SU-8 thickness was carried out by coating SU-8 on Au/Cr/glass-slide with one end Au exposed. The test was set up by soaking one end of the slides (a single layer of SU-8 covered) in electrolyte (50 mM PB with 100 mM K2SO4, pH 7), and measuring the leakage current from another end of the slide via a counter electrode which was also immersed in the electrolyte. The results showed a leakage current lower than 1 nA for the 20 lm thick SU-8, and lower than 1 pA for the 40 lm thick SU-8 after 30 min soaking period. A 40 lm thick of SU-8 was therefore chosen for the passivation layer of the sensing linear array. The adhesion between the SU-8 and the substrate was improved by using a double layer processing including a thin layer plus a thick layer. First, a 2 lm thick SU-8-2 was patterned and cured, and then a 40–50 lm thick of SU-8 50 was spin coated at 3000 rpm speed, and soft baked on a hotplate at 65 °C for 5 min and 95 °C for 15 min. After exposure, the wafer was post baked on a hot plate and developed in EC Solvent for 6 min. Optical inspection of the patterned SU-8 layer revealed ‘‘crack” like features on the corners of the sensing pads. However, SEM inspection (Fig. 3) showed ‘‘wrinkles” on the surface. The wafer was diced and cleaned, and a MOSFET chip was mounted on each glass die, as shown in Fig. 4, followed by mounting onto a PCB board. Prior to wire bonding of the sensing gold pads to MOSFET chip, the transistor source, drain and gate contacts were grounded to avoid any damage to the gates due to electrostatic discharge during wire bonding. After wire bonding, all the bonding wires, bond pads and MOSFET chip were encapsulated by epoxy-based Glob-Top. 3. Test results A peptide aptamer prepared by insertion of the peptide pep2 into the scaffold protein STM (Stefin A Triple Mutant), was used for the protein CDK2 detection. Before testing, the gold sensing linear array was thoroughly cleaned in an UV-ozone system followed by ethanol and DI water rinse to remove any impurities on the gold surface. This surface modification is critical in obtaining a biocompatible surface for the bimolecular attachment, test sensitivity and stability. The devices’ characteristics were tested before protein interaction to make sure there was no hysteresis on the sensing surface. The testing process included the STM-pep2 immobilization (1 lM solution in 10 mM phosphate buffer) onto the clean gold sensing pads, followed by passivation of any exposed gold areas
JSR photoresist
LOR (lift-off resist )
Fig. 2. SEM cross-section of photoresist with undercut profile for gold lift-off process.
Fig. 3. A gold sensing pad with SU-8 on; the inset is a SEM cross-section image shows one corner of the SU-8 layer.
755
L. Wang et al. / Microelectronic Engineering 87 (2010) 753–755
1m
MOSFET chip
100µ 10µ
Gold sensing pads
-ID (A)
5mm
1µ 100n
+15mV
10n
Fig. 4. Optic picture of external MOSFET chip and gold sensing linear array on glass substrate.
1n 100p
STM-pep2 rCDK2 JB-4 / P1 VDS=-1V
-1.00
-0.75
-0.50
-0.25
VGS (V) Fig. 6. Transfer characteristics of a STMCys+-pep2 interaction with CDK2 on gold extended-gate MOSFET, a p-type transistor with gate length 0.7 lm and width 3860 lm.
bilized peptide aptamer STM-pep2, a shift of 15 mV in the transfer characteristic is observed. The variation of the signal with the concentration of protein was measured by injecting known amounts of purified CDK2 into the cell containing the functionalized gold electrodes. The obtained binding curve follows a logistic model, which is characteristic of immunoassays. Further testing results [6] showed that shifts of the order of 1 mV can be reproducibly detected with CDK2 concentrations much lower than 60 nM indicated in the initial test result. To increase the test sensitivity further, a new design with decreased sensing pad size is currently being fabricated. Fig. 5. No shift on the gate voltage is observed (in this case for a p-type FET) when there is no target protein introduced.
by a thiolated poly ethylene glycol in order to reduce non-specific interactions [6]. During testing, the substrate of the MOSFETs was grounded through the source electrode, with the drain to source voltage VDS kept at 1 V and 1 V for n-type and p-type transistors, respectively. The gate to source voltage VGS was applied through an Hg/Hg2SO4 reference electrode immersed in the electrolyte, and varied from 0.5 V to +2.0 V. The transfer characteristics of the BioFET were measured before and after a target protein CDK2 was interacted with the peptide aptamer immobilized on the sensing pads. Upon biomolecular interaction, variations in the charge distribution of the bio-layer occur, which in turn induce variations in the potential at the interface between the electrolyte and the extended gold sensing pad. Since the gate voltage is applied via the reference electrode immersed in the electrolyte, a parallel shift on the transfer characteristics of the device is expected when a protein interaction induces a change in the bio-layer charge distribution. Fig. 5 shows the transfer characteristics before and after a biological sample (yeast lysate) that does not contain the target protein is interacted with the system. Since no interaction occurs, no shift on the transfer characteristic is observed. An initial test result of the biomolecular interaction is shown in Fig. 6. When a sample containing 60 nM of target CDK2 interacts with the immo-
4. Conclusion A linear array of BioFET was fabricated using extended gold gates connecting to MOSFETs as a convenient sensing approach. The fabrication of the gold sensing pads and SU-8 passivation layer were optimized to meet the protein detection requirements. The test results showed the applicability of the linear array BioFET sensing detection on the cancer cell marker proteins by using peptide aptamers detection system. Acknowledgments This work is supported by the Biotechnology and Biological Science Research Council (BBSRC) UK under the Grant BB/D523094/ 01. References [1] [2] [3] [4] [5]
P. Bergveld, IEEE Trans. Biomed. Eng. 17 (1970) 70–71. P. Bergveld, Sens. Actuator A 56 (1996) 65–73. M.J. Schöning, Sensors 5 (2005) 126–138. P. Estrela, P. Migliorato, J. Mater. Chem. 17 (2007) 219–224. M. Crawford, R. Woodman, P. Ko Ferrigno, Brief. Funct. Genomics Proteomics 2 (1) (2003) 72–79. [6] P. Estrela, D. Paul, Q. Song, L.K.J. Stadler, L. Wang, E. Huq, J.J. Davis, P. Ko Ferrigno, P. Migliorato, submitted for publication.
Contents lists available at ScienceDirect
Microelectronic Engineering journal homepage: www.elsevier.com/locate/mee
Fabrication of BioFET linear array for detection of protein interactions Ling Wang a,*, Pedro Estrela b,1, Ejaz Huq a, Peng Li b, Stephen Thomas a, Paul Ko Ferrigno c, Debjani Paul b, Paul Adkin a, Piero Migliorato b a
Technology Department, Rutherford Appleton Laboratory, STFC, Oxfordshire, OX11 0QX, UK Department of Engineering, University of Cambridge, Cambridge CB3 0FA, UK c Leeds Institute of Molecular Medicine, St. James’ University Hospital, Leeds LS9 7TF, UK b
a r t i c l e
i n f o
Article history: Received 14 September 2009 Received in revised form 26 November 2009 Accepted 26 November 2009 Available online 2 December 2009 Keywords: BioFETs Protein interaction SU-8
a b s t r a c t An extended-gate MOSFETs (metal–oxide-semiconductor field-effect transistors) based biosensing linear array has been fabricated for label-free protein interaction detection. The device was realized using a combination of very low leakage current MOSFET transistors and an external gate where the chemical reaction would take place. Peptide aptamers that recognize cyclin-dependent kinase (CDK), a protein cancer marker, were used as a biological test system. The test results showed a high sensitive in the detection of CDK. Ó 2009 Elsevier B.V. All rights reserved.
1. Introduction There has been a growing interest in silicon-based field-effect biosensors since Bergveld first investigated the use of ion-sensitivity field-effect transistors (ISFETs) three decades ago [1,2]. Biologically active field-effect transistors (BioFETs) are miniaturized silicon-based electrochemical biosensors that can detect biomolecular interactions occurring at the gates of MOSFETs. Compared to the conventional optical and mass spectrometry detection methods, BioFETs provide fast, label-free, high sensitivity and direct electronic signal readout [3,4]. FET-based biosensors have the potential for developing high-throughput label-free microarrays for applications in clinical diagnostics, drug discovery, food monitoring, and bioterrorist security inspection. In this paper we present a convenient approach for the fabrication of a linear array extended-gate BioFETs. MOSFETs were used as electrical transducers with the gates connected to external gold sensing pads where the electrolyte is introduced and biomolecular interactions take place. This approach allows an easy separation between the sensing area and the transistor in order to reduce any risk of leakage. It also enables an easy way to define and modify the sensing area and shape, and avoiding post-processing on MOSFET itself. The operation of the BioFET detection is based on the immobilization of probe molecules on the extended gate, * Corresponding author. Tel.: +44 1235 445829; fax: +44 1235 446283. E-mail address: [email protected] (L. Wang). 1 Present address: Department of Electronic and Electrical Engineering, University of Bath, Bath BA2 7AY, UK. 0167-9317/$ - see front matter Ó 2009 Elsevier B.V. All rights reserved. doi:10.1016/j.mee.2009.11.148
which is placed in contact with the electrolyte. The gate voltage is applied via a reference electrode immersed in the electrolyte (see Fig. 1). Upon a biomolecular interaction, the change in charge distribution of the bio-layer induces a variation in the potential drop at the interface with the electrolyte, which in turn modulates the gate voltage. Peptide aptamers are artificial recognition molecules made from an insertion of a variable peptide sequence into a constant scaffold protein [5], which allow specific protein targets to bind and get recognized. With the advantages of excellent recognition specificity and high binding affinity compared to antibodies, the peptide aptamers have many potential applications in protein detections. In this paper, the interaction of different peptide aptamers with CDK2 proteins was detected by measuring the transfer characteristics of the MOSFETs.
2. Fabrication The MOSFETs were fabricated at IMEC (Belgium), with a number of n-type and p-type transistors on each chip. The gate lengths of the MOSFETs ranged from 0.7 lm to 2.0 lm, and widths ranged from 3860 lm to 1340 lm with a constant gate area. The MOSFETs were fabricated with 17 nm thick of gate oxide. The operation voltage is 3.3–5 V with a saturation current of 358 lA/lm. There are no electrostatic discharge protection elements in the MOSFET devices in order to minimize the leakage current. The leakage current is a critical concern in this type of detection, since a very small charge change during protein interaction will be detected. A calculation
754
L. Wang et al. / Microelectronic Engineering 87 (2010) 753–755
Reference Electrode
VGS
VDS
Extended gate
n+
n+
Source
Drain P-type Si
Fig. 1. Schematic structure of BioFET detection of protein.
result showed that the gate leakage current must be lower than 10 14 A. The linear array sensing pads were designed 500 500 lm gold pads separated by 1 mm for ease of fabrication. The gold sensing pads were fabricated on glass substrates using a lift-off process. A double layer of photoresist process was used for high quality gold pad edge profile, especially for the very long and thin gold lines, which connect the sensing pads and the wire-bond pads. A layer of LOR (lift-off resist) 20B was spin coated and baked at 180 °C on a hot plate for 10 min. followed by a layer of JSR photoresist spin coated and baked, which resulting in a 2 lm thick LOR and 1.2 lm thick JSR resist. After exposure on a Karl Suss Mask Aligner [UV400 light] the JSR photoresist was developed and hard baked. The LOR was then etched in dilute AZ400 K for about 2 min. to obtain an undercut profile. An SEM picture of the cross-section is shown in Fig. 2. The gold layer (100 nm thick) was deposited in an electron beam evaporator (SV2000) with a thin layer of chrome (10 nm thick) as seed layer. In order to obtain a very smooth gold surface which is required for the protein binding, the gold film was deposited with a low deposition rate of 0.5 Å/s. The lift-off process was achieved in acetone followed by soaking in dilute AZ400 K to remove the LOR. SU-8 epoxy-based photoresist was used as passivation layer to cover the whole surface except the sensing pads and wire-bond
pads. The advantages of the SU-8 include its better biocompatibility compared to silicon dioxide, easier to pattern and good insulation properties. The disadvantage of SU-8 is its weak adhesion to glass surface and stress problem. An investigation for the SU-8 process was performed to optimize the thickness, exposure dose and baking conditions. A series of tests for optimization of SU-8 thickness was carried out by coating SU-8 on Au/Cr/glass-slide with one end Au exposed. The test was set up by soaking one end of the slides (a single layer of SU-8 covered) in electrolyte (50 mM PB with 100 mM K2SO4, pH 7), and measuring the leakage current from another end of the slide via a counter electrode which was also immersed in the electrolyte. The results showed a leakage current lower than 1 nA for the 20 lm thick SU-8, and lower than 1 pA for the 40 lm thick SU-8 after 30 min soaking period. A 40 lm thick of SU-8 was therefore chosen for the passivation layer of the sensing linear array. The adhesion between the SU-8 and the substrate was improved by using a double layer processing including a thin layer plus a thick layer. First, a 2 lm thick SU-8-2 was patterned and cured, and then a 40–50 lm thick of SU-8 50 was spin coated at 3000 rpm speed, and soft baked on a hotplate at 65 °C for 5 min and 95 °C for 15 min. After exposure, the wafer was post baked on a hot plate and developed in EC Solvent for 6 min. Optical inspection of the patterned SU-8 layer revealed ‘‘crack” like features on the corners of the sensing pads. However, SEM inspection (Fig. 3) showed ‘‘wrinkles” on the surface. The wafer was diced and cleaned, and a MOSFET chip was mounted on each glass die, as shown in Fig. 4, followed by mounting onto a PCB board. Prior to wire bonding of the sensing gold pads to MOSFET chip, the transistor source, drain and gate contacts were grounded to avoid any damage to the gates due to electrostatic discharge during wire bonding. After wire bonding, all the bonding wires, bond pads and MOSFET chip were encapsulated by epoxy-based Glob-Top. 3. Test results A peptide aptamer prepared by insertion of the peptide pep2 into the scaffold protein STM (Stefin A Triple Mutant), was used for the protein CDK2 detection. Before testing, the gold sensing linear array was thoroughly cleaned in an UV-ozone system followed by ethanol and DI water rinse to remove any impurities on the gold surface. This surface modification is critical in obtaining a biocompatible surface for the bimolecular attachment, test sensitivity and stability. The devices’ characteristics were tested before protein interaction to make sure there was no hysteresis on the sensing surface. The testing process included the STM-pep2 immobilization (1 lM solution in 10 mM phosphate buffer) onto the clean gold sensing pads, followed by passivation of any exposed gold areas
JSR photoresist
LOR (lift-off resist )
Fig. 2. SEM cross-section of photoresist with undercut profile for gold lift-off process.
Fig. 3. A gold sensing pad with SU-8 on; the inset is a SEM cross-section image shows one corner of the SU-8 layer.
755
L. Wang et al. / Microelectronic Engineering 87 (2010) 753–755
1m
MOSFET chip
100µ 10µ
Gold sensing pads
-ID (A)
5mm
1µ 100n
+15mV
10n
Fig. 4. Optic picture of external MOSFET chip and gold sensing linear array on glass substrate.
1n 100p
STM-pep2 rCDK2 JB-4 / P1 VDS=-1V
-1.00
-0.75
-0.50
-0.25
VGS (V) Fig. 6. Transfer characteristics of a STMCys+-pep2 interaction with CDK2 on gold extended-gate MOSFET, a p-type transistor with gate length 0.7 lm and width 3860 lm.
bilized peptide aptamer STM-pep2, a shift of 15 mV in the transfer characteristic is observed. The variation of the signal with the concentration of protein was measured by injecting known amounts of purified CDK2 into the cell containing the functionalized gold electrodes. The obtained binding curve follows a logistic model, which is characteristic of immunoassays. Further testing results [6] showed that shifts of the order of 1 mV can be reproducibly detected with CDK2 concentrations much lower than 60 nM indicated in the initial test result. To increase the test sensitivity further, a new design with decreased sensing pad size is currently being fabricated. Fig. 5. No shift on the gate voltage is observed (in this case for a p-type FET) when there is no target protein introduced.
by a thiolated poly ethylene glycol in order to reduce non-specific interactions [6]. During testing, the substrate of the MOSFETs was grounded through the source electrode, with the drain to source voltage VDS kept at 1 V and 1 V for n-type and p-type transistors, respectively. The gate to source voltage VGS was applied through an Hg/Hg2SO4 reference electrode immersed in the electrolyte, and varied from 0.5 V to +2.0 V. The transfer characteristics of the BioFET were measured before and after a target protein CDK2 was interacted with the peptide aptamer immobilized on the sensing pads. Upon biomolecular interaction, variations in the charge distribution of the bio-layer occur, which in turn induce variations in the potential at the interface between the electrolyte and the extended gold sensing pad. Since the gate voltage is applied via the reference electrode immersed in the electrolyte, a parallel shift on the transfer characteristics of the device is expected when a protein interaction induces a change in the bio-layer charge distribution. Fig. 5 shows the transfer characteristics before and after a biological sample (yeast lysate) that does not contain the target protein is interacted with the system. Since no interaction occurs, no shift on the transfer characteristic is observed. An initial test result of the biomolecular interaction is shown in Fig. 6. When a sample containing 60 nM of target CDK2 interacts with the immo-
4. Conclusion A linear array of BioFET was fabricated using extended gold gates connecting to MOSFETs as a convenient sensing approach. The fabrication of the gold sensing pads and SU-8 passivation layer were optimized to meet the protein detection requirements. The test results showed the applicability of the linear array BioFET sensing detection on the cancer cell marker proteins by using peptide aptamers detection system. Acknowledgments This work is supported by the Biotechnology and Biological Science Research Council (BBSRC) UK under the Grant BB/D523094/ 01. References [1] [2] [3] [4] [5]
P. Bergveld, IEEE Trans. Biomed. Eng. 17 (1970) 70–71. P. Bergveld, Sens. Actuator A 56 (1996) 65–73. M.J. Schöning, Sensors 5 (2005) 126–138. P. Estrela, P. Migliorato, J. Mater. Chem. 17 (2007) 219–224. M. Crawford, R. Woodman, P. Ko Ferrigno, Brief. Funct. Genomics Proteomics 2 (1) (2003) 72–79. [6] P. Estrela, D. Paul, Q. Song, L.K.J. Stadler, L. Wang, E. Huq, J.J. Davis, P. Ko Ferrigno, P. Migliorato, submitted for publication.