www.ietdl.org Published in IET Nanobiotechnology Received on 25th September 2013 Revised on 9th October 2013 Accepted on 11th October 2013 doi: 10.1049/iet-nbt.2013.0061
ISSN 1751-8741
Design and fabrication of field-effect biosensors for biochemical detection Shoucai Yuan, Xiaolin Fan, Ziyu Wang School of Physics and Electronics Information, GanNan Normal University, Economic and High-Tech. Development Zone, Ganzhou, JiangXi 341000, People’s Republic of China E-mail: [email protected]
Abstract: Biochemically sensitive field-effect sensors are fabricated with simplified chip technology. Its fabrication process flow is designed based on metal gate complementary metal-oxide semiconductor technology, in which only six pattern masks are employed. The sensors are measured as field modulation resistors since they are made in its depletion mode. The milliampere magnitude response of conducting currents from certain biochemical materials achieves distinct sensitivity when measured on our fabricated sensors with different sensitive areas of W/L = 4.2 and 20.0. To check the stability of the sensor, up to 20 repeated tests are conducted on the same sensor chip operated in its three states, in which no materials (blank state, called ‘blank’), pure water and biochemical materials are coated on its gate dielectric film, respectively. Measured results show that the response currents for certain materials are distributed in certain current range. Taking the response current of blank as a reference value, the response current of pure water is positive but very close to that of blank because of the small electric dipole properties of pure water. However, the response current of biochemical materials are negative and far apart from that of blank, because the biochemical materials have large electric dipole properties and clearly show measurement resolution.
1
Introduction
Chemically sensitive field-effect transistors (ChemFETs) [1–3] have proved their potential as microsensors for biomedical analytics, since they show great advantages over conventional ion-selective electrodes in terms of small dimensions, low-output impedance, fast response, mass-fabrication ability and possibility for integration in smart sensor arrays [4–6]. Among ChemFET devices, the ion-sensitive field-effect transistors (ISFETs) are best-known. However, the practical applications have not been widely found since the ISFET concept was proposed by P. Bergveld in 1970 [7, 8]. The main problem is the drift time [9, 10] of the sensor response, which brings a strong limitation mainly to the determination of ion activities and concentrations in the physiological range (e.g. blood electrolytes). A further limiting factor is the relatively high manufacturing cost of those sensitive devices. At present, there is a trend to choose single-use devices which are made by very cheap materials and simple fabrication technologies in biomedical applications. In this paper, we employ a simplified silicon chip fabrication technology to make ChemFET serve as biochemical sensors (biosensors) with low cost. We also use self-assembled molecules (SAM) [11, 12] techniques to coat analyte polymer/transducer on the top of gate dielectric film instead of gate metal of the conventional metal-oxide semiconductor field-effect transistor (MOSFET). Aspects of this work have already been published in a recent conference [13]. However, we conduct wide experiments and extract detailed data results in this new paper. Our 208 & The Institution of Engineering and Technology 2014
biosensors overcome the drift time problem as well, because of the field-enhanced conductivity. Therefore, they will be the most popular low-cost, single-use sensors which can detect not only polar molecules [14–20] but also ions, neutral molecules and biochemical materials, as long as appropriate transducers (recognition layer) are added to the gate dielectric film to provide specificity.
2 Field-effect biosensor fabrication and its sensitive characteristics The biosensor we proposed is a field-effect transistor in which the metallic gate is omitted so that the dielectric material can be exposed to the analyte polymer directly from solution or via a transducer layer with a specific recognition function. The cross-sections of the conventional MOSFET and ChemFET are illustrated comparatively in Figs. 1a and b, respectively. In the conventional MOSFET, the p-type silicon wafer as substrate forms the devices channel, whereas the dielectric film and the aluminium layer form the gate electrode, which are all protected by passivation films. In the ChemFET structure, the aluminium layer which covered the gate dielectric film is not required because the dielectric layer is responsible for the biochemical sensing. Therefore the passivation film is also completely removed in a region delimitated by the area of the source and drain electrodes. To fabricate a ChemFET sensor in Fig. 1b, the process flow is illustrated in Fig. 2, the p-type silicon wafer with a IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
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Fig. 1 Comparative schematic representation of conventional MOSFET and ChemFET a MOSFET with gate metal b ChemFET without gate metal
Fig. 2 Fabrication processes flow of our designed ChemFET sensor (biosensor) IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
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www.ietdl.org thickness of 525 μm is used as a substrate and performs as an active channel in the device operation. First, mask 1# is patterned, by means of conventional lithography and dry etching, to define the p+ guard ring for substrate grounding. Boron (B) is implanted with energy: 40 keV, dosage: 1.0 × 1015 cm−2 and concentration: 1.0 × 1019 cm−3 and the measured sheet resistance is 68 Ω/□. Second, mask 2# is patterned to define the devices drain and source area. Arsenic (As) is implanted with energy: 100 keV, dosage: 4.0 × 1015 cm−2 and concentration: 1.0 × 1019 cm−3 and the measured sheet resistance is 48 Ω/□. Subsequently, mask 3# is patterned to define the devices active area. The 104.3 nm silicon oxide (SiO2) and 18.6 nm silicon nitride (Si3N4) are deposited on the active area to form two layers of complex gate dielectric films. The Si3N4 layer performs a double role in this case: it represents the dielectric film of sensitive gate of biosensors, as well as a stopping layer for the passivation film etching. The thickness of these two layers of gate dielectric film has been chosen to enhance the detection sensitivities. Threshold voltage adjusting is also achieved in this step. Phosphorus (P) is implanted with energy: 150 keV, dosage: 1.5–2.0 × 1013 cm−2 and concentration: 6.5 × 1015 cm−3, thus, the conduction channel occurs at the semiconductor surface under the gate dielectric film. Then, mask 4# is patterned to define the devices Ohmic contact area for drain, source and p-substrate grounding, and mask 5# is patterned to define the metal aluminium connection. Finally, mask 6# is patterned to define the passivation film and open the biosensor’s sensitive window, where the metal and passivation film are all removed from the top of the gate dielectric film. The fabrication process parameters of our design and measurement are listed in Table 1 and a microphotograph of our fabricated biosensor chip is shown in Fig. 3. After cleavage, the die is wire-bonded and encapsulated with a two-component epoxy called EP42HT at a temperature of 300 K or so.
Fig. 3 Microphotograph of our fabricated biosensor chips (portion)
Fig. 4 Schematic diagram showing the bias conditions used for testing the biosensors (ChemFET sensors) with Ag/AgCl reference electrodes
Table 1 Design and measurement processes parameters for our fabricated biosensors No.
Processes
1
boron (B) implantation
2
drain/source implantation
3
drive in
4
phosphorus (P) implantation
5
gate oxidation (SiO2)
6 7
silicon nitride (Si3N4) contact low pressure chemical vapour deposition (LPCVD) (SiO2) passivation LPCVD (Si3N4) atmospheric pressure (AP) CVD (SiO2) APCVD (SiO2) Al/Si Al/Cu
8 9 10 11 12
Controlled parameters 40 keV, 1.0 × 1015 cm−2 100 keV, 4.0 × 1015 cm−2 xj = 3.5 μm 150 keV, 1.5– 2.0 × 1013 cm−2 100 ± 15 nm
Measurement parameters sheet resistance :68 Ω/□ sheet resistance: 48 Ω/□ SiO2 on p+ 59 nm, on n+ 89 nm concentration: 6.5 × 1015 cm−3
20 ± 10 nm 175 ± 17 nm
104.3 nm, Qss 1 × 1010 C/cm2 18.6 nm 189.4 nm
60 ± 6 nm
60 nm
720 ± 72 nm
805.6 nm
175 ± 25 nm 0.4 ± 0.2 μm 1.6 ± 0.2 nm
188.1 nm 0.38 μm 1.5 nm
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Fig. 5 Repeatability and stability testing for our biosensors with detection materials of pure water and S. cercariae protein a Chip 1# b Chip 2# IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
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Fig. 6 Current response related with number of carboxyl group of serial alkyl-acids and different functional groups of serial organic polar acids a Carboxyl group varied for serial alkyl-acids (chip with W/L = 4.20) b Functional group varied for serial organic polar acids (two chips of W/L = 4.20 and W/L = 20.0)
To measure the current response of the detected materials, the terminal bias conditions are illustrated in Fig. 4. The biosensors and an Ag/AgCl reference microelectrode forming a three-terminal device are biased in the same way as a conventional MOSFET in the common source configuration, in which VGS (V) is gate–source bias voltage and VDS (V) is drain–source bias voltage. The output drain– source current, IDS (mA), which indicates the current response for certain detected biomolecules and acids, is recorded by a HP 4156A Precision Semiconductor Parameter Analyzer at a temperature of 300 K. Despite the fact that the conduction mechanisms in biosensors are not clearly elucidated [7, 21–23], their ‘macroscopic’ electronic behaviour may be reasonably described by the same equations as for conventional MOSFETs. At low gate bias voltages, the channel carrier mobility can be assumed as a constant and the current (IDS) flowing through the channel between source and IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
drain is well described by the equation [24, 25]
IDS
2 W VDS = · m · Cins · VGS − Vth VFB · VDS − 2 L W + · s · dsemi · VDS L
(1)
where μ is the channel carriers mobility and σ is the bulk conductivity, Cins is the dielectric capacitance per unit area; dsemi is the thickness of the semiconductor surface depletion layer, VFB is the flat-band voltage and Vth (VFB) is the threshold voltage which is the function of VFB. Therefore the biosensors’ current response of serial polar acids and certain biochemical materials can be attributed to the change in the flat-band voltage induced by the variation of 211
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www.ietdl.org the potential at the electrolyte/gate dielectric interface [26] VFB = Eref + xSOL −
wsemi Qins − − c0 q Cins
(2)
where Eref is the potential of the Ag/AgCl reference electrode, χ SOL is the surface dipole potential of the detected materials, wsemi/q is the semiconductor work function and Qins is the effective charge per unit area and c0 is the voltage drop at the electrolyte/dielectric interface. As shown in (1) and (2), for given devices bias voltage of VDS, VGS and sensitive area parameters of W/L, the devices’ output drain–source current is linearly correlated to χ SOL, c0 (parameters of detected materials) and Qins/Cins, wsemi/q etc. (parameters of fabricated biosensors). In our experiment, biochemical materials combined selected transducers, certain biomolecules and acids can serve as the analyte polymer which is coated on a sensitive area of the biosensor instead of the gate metal by self-assembly techniques (SAM). Thus, work function and charge distribution in gate dielectric film will be changed. Although there is no direct interaction between the semiconductor surface and the analyte polymer, the threshold voltage (Vth) related to semiconductor work function (wsemi/q) is varied. Moreover, because the presence of the analyte polymer has modified the electric field in the surface of the channel, the channel conductivity (σ) is changed accordingly. The change of Vth and σ will vary the biosensor’s output current (IDS) accordingly, which forms the current response of the detected materials. The calculated data of this simple model and the experiments data will be compared in the following section.
Fig. 7 Current response sensitivity related with biosensors sensitive areas (W/L) and current response comparison between simple analytical model calculation and measurement according to variation of haemocyanin concentration a Current response sensitivity related with biosensors sensitive areas (W/L) b Current response comparison between simple analytical model calculation and measurement according to variation of haemocyanin concentration (W/L = 20.0) 212 & The Institution of Engineering and Technology 2014
3 Electrical characterisation results and discussions To optimise the biosensor’s sensitive area, series W/L channel parameters are designed on the same die and fabricated by the same technology. In addition, the devices are made in their depletion operation mode, that is, the devices’ channel is already implanted with an n-type doping agent. Hence, it is always in a weak conducted state and a current will flow through its channel between the drain and the source electrodes even if no gate voltage (VGS = 0.0 V) is applied. Therefore it will show typical resistor output characteristics when VGS = 0.0 V bias conditions are used in Fig. 4. First, we conduct the repeatability and stability test for our fabricated biosensors, by using the measurement schematic diagram shown in Fig. 4, with device parameters of Vth = −1.0 V and device terminal bias conditions of VGS = 0.0 V, VDS = 3.0 V. We selected two biosensor chips with the same sensitive areas of W/L = 20.0 marked as chip 1# and chip 2#, respectively. Then, pure water and protein extracted from Schistosoma cercariae are used as detected materials.
Fig. 8 Current response distribution and comparison among the blank, pure water and S. cercariae worms with two chips a W/L = 4.2 b W/L = 20.0 IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
www.ietdl.org The biosensor’s drain–source response currents for each detected material are measured 10–20 times, in order to verify the repeatability and stability performances of our fabricated biosensor. The process for our measurement is, firstly, coating the detected materials, such as, pure water or protein of S. cercariae on top of the sensor’s gate dielectric film as the analyte polymer with the same SAM formation conditions, and then the sensor’s drain–source current is measured. As shown in Fig. 5, it is clearly demonstrated that the detection experiment is not stable in its initial 1–10 times measurements; however, after this initial period of time, the later detection experiment shows a relatively stable response current on our fabricated biosensor. Then, the detection experiments are conducted for series acids, for example, hexanoic acid, adipic acid and citric acid and so on. We have checked and compared the biosensor’s detection characteristics according to the carboxyl group and functional group of selected acids. Fig. 6a show the measured biosensor’s drain–source current varying with the acids carboxyl group. The blank state measurement data are marked ‘blank’ in the figure. It can be seen that the response current decreases with the increase of carboxyl group for serial alkyl-acid, since there are 1, 2 and 3-carboxyl group for hexanoic, adipic and citric acids, respectively. Those current responses are all smaller than the current response of the blank state. Fig. 6b gives the relationship among detected response current and the functional group of serial polar acid for two sensitive areas of W/L = 4.20 and 20.0. This current response decreases with the order of blank, –OOH, –OH and –NH2 for both sensitive areas of W/L = 4.20 and 20.0. The biosensor’s quantified sensitivity [27] is measured according to the series sensitive area and different
concentrations of target detected biomolecules and acids. Haemocyanin, the protein extracted from snails, is used in this experiment. Fig. 7a shows the current response of haemocyanin for series W/L values, here the sensitivity is defined as response current per W/L values, that is, sensitivity = IDS/(W/L). It is clear that the sensitivity decreases with the increase of W/L. Fig. 7b demonstrates the current response with different haemocyanin concentrations from 0.1753 to 0.7109 (mg/ml) for sensitive area of W/L = 20.0 chips. The measurement data are also compared with the calculated data from the simple model represented by formula (1), in which the main model parameters used in our calculation are listed as W/L = 20, VDS = 3 (V), μn = 450 (cm2/V/s), dsemi = 30 (nm), gate oxide film (SiO2): tox = 104.3 (nm), εr = 3.9, gate silicon nitride film (Si3N4): tox = 18.6 (nm), εr = 6.2, Nsub = 1.7 × 1016 (cm−3), Vth = −1.0 (V) and σ = q·nsub·μn (Siemens/cm). We assume that, for each measured target molecule concentration, Vth and σ will vary and its value will be determined by measurement and calculation. Fig. 7b shows that the higher the haemocyanin concentration, the larger response current we can detect, and those response currents are all larger than that of blank. The next detection experiment is conducted for S. cercariae worms [28] with different sensitive area, the measurement results are compared with that of pure water and blank, as shown in Fig. 8, where (a) is for W/L = 4.20 and (b) is for W/L = 20.0. It can be seen that the response current for pure water and blank is relatively stable under certain experimental conditions. However, the response currents for S. cercariae worms are distributed in a certain current range. The reason for this data distribution range is (i) the variety of density for S. cercariae worms, (ii) the
Fig. 9 Current response for detection of S. cercariae worms using linoleic acids as transducers (recognition) layer a W/L = 4.20 b W/L = 20.0 IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
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www.ietdl.org different positions on top of the sensitive area for each S. cercariae worms after finishing coating on the dielectric film, (iii) the characteristics of SAM S. cercariae worms and its formation conditions is related to each SAM processing. Since there are no or little electric dipole characteristics in pure water, the response current of pure water is positive and close to that of blank. However, S. cercariae worms have high dipole charge characteristics, properties of unique electronic conduction and delocalised electronic structure along the conjugated backbone, and its measured response currents are negative and far apart from that of pure water and blank. In all the above experiments, the effect of surface charging state density and especially pH value of pure water to the biosensor’s response current must be considered in our experiment for blank and pure water test. We assume that the influence of surface charging state density is already consistent in the output current response. However, water’s pH value has large influence on the surface charging state density, the surface charging state density will increase with the increase of pH value [29]. Hence, we must ensure the pH value to be stable during our experiment, that is, using the same barrel/batch of pure water with pH value of 6.8 ± 0.2. Otherwise, our detection results comparison will be invalid since the pure water pH value is varied with temperature and environment. For example, the research results in [29] show that the surface charging state density varies rapidly when the pH value varies from 5 to 7, which will make the measurement data unstable and the experiment uncontrollable. The research results in [29] also show that, the surface charging state density changes slowly when the pH value is smaller than 5. Hence, serial polar acids can be used as transducers [1, 30] to detect S. cercariae worms. For example, we use linoleic acid as transducer materials in one of our experiments. Fig. 9 shows our experiment in repeated detections of S. cercariae worms by using linoleic acid as a transducer film. The test processes are repeatedly carried out seven times. For each time of measurement, we first make linoleic acid transducers using SAM processes, and measure the sensor response current for this transducer. Then, the S. cercariae worms are coated on top of the pervious transducer film also by SAM processes. The synthesis-induced response current is measured and compared with the current induced only by the transducer film. As shown in Fig. 9, (a) is for W/L = 4.20 and (b) for W/L = 20.0. Generally, after the SAM S. cercariae worms, the synthesis-induced response currents are larger than that of linoleic acid transducers except for the second time measurement in Fig. 9b.
4
Conclusions
We have designed and fabricated ChemFET sensors (biosensors) on p-type silicon substrates with channel parameters of W/L = 4.2, 20.0 and so on. We have also investigated the ability of our fabricated biosensors to perform charge detection in aqueous media. We conduct the repeatable and stable experiments 10–20 times, using pure water and S. cercariae protein as detected materials and compared their response current results. The detecting ability for series acids is verified by comparing the sensor’s current response for specifically selected acid with different carboxyl group and functional group. The detected results are also compared with that of blank. The biosensor’s sensitivity is verified by checking the variation of the 214 & The Institution of Engineering and Technology 2014
current response with serial sensitive area (W/L) and serial haemocyanin concentrations, which is compared with simple model calculation data. After finishing the above experiment, we selected two groups of sensors for S. cercariae worms’ detection, and the detected response currents are compared with that of pure water and blank. For sensitive areas of W/L = 4.2 and 20.0, the current range for S. cercariae worms are 0.145–0.060 mA and 1.650– 1.300 mA, respectively. The minimum current differences (resolution spacing) compared with blank were 0.075 and 0.350 mA, respectively. The transducer (recognition) layer experiment is also conducted using linoleic acid as transducer for detecting S. cercariae worms, which gives clearly detected resolution in our experiment.
5
Acknowledgments
The authors would like to thank Prof. and PhD J.Q. Han and senior engineer Mr. X.L. Wang in the Microelectronics Institute for their support of wafer fabrication and valuable feedback. This research is supported by the Natural Science Fund Committee of China (NSFC) (no. 51377025, GJJ13661).
6
References
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ISSN 1751-8741
Design and fabrication of field-effect biosensors for biochemical detection Shoucai Yuan, Xiaolin Fan, Ziyu Wang School of Physics and Electronics Information, GanNan Normal University, Economic and High-Tech. Development Zone, Ganzhou, JiangXi 341000, People’s Republic of China E-mail: [email protected]
Abstract: Biochemically sensitive field-effect sensors are fabricated with simplified chip technology. Its fabrication process flow is designed based on metal gate complementary metal-oxide semiconductor technology, in which only six pattern masks are employed. The sensors are measured as field modulation resistors since they are made in its depletion mode. The milliampere magnitude response of conducting currents from certain biochemical materials achieves distinct sensitivity when measured on our fabricated sensors with different sensitive areas of W/L = 4.2 and 20.0. To check the stability of the sensor, up to 20 repeated tests are conducted on the same sensor chip operated in its three states, in which no materials (blank state, called ‘blank’), pure water and biochemical materials are coated on its gate dielectric film, respectively. Measured results show that the response currents for certain materials are distributed in certain current range. Taking the response current of blank as a reference value, the response current of pure water is positive but very close to that of blank because of the small electric dipole properties of pure water. However, the response current of biochemical materials are negative and far apart from that of blank, because the biochemical materials have large electric dipole properties and clearly show measurement resolution.
1
Introduction
Chemically sensitive field-effect transistors (ChemFETs) [1–3] have proved their potential as microsensors for biomedical analytics, since they show great advantages over conventional ion-selective electrodes in terms of small dimensions, low-output impedance, fast response, mass-fabrication ability and possibility for integration in smart sensor arrays [4–6]. Among ChemFET devices, the ion-sensitive field-effect transistors (ISFETs) are best-known. However, the practical applications have not been widely found since the ISFET concept was proposed by P. Bergveld in 1970 [7, 8]. The main problem is the drift time [9, 10] of the sensor response, which brings a strong limitation mainly to the determination of ion activities and concentrations in the physiological range (e.g. blood electrolytes). A further limiting factor is the relatively high manufacturing cost of those sensitive devices. At present, there is a trend to choose single-use devices which are made by very cheap materials and simple fabrication technologies in biomedical applications. In this paper, we employ a simplified silicon chip fabrication technology to make ChemFET serve as biochemical sensors (biosensors) with low cost. We also use self-assembled molecules (SAM) [11, 12] techniques to coat analyte polymer/transducer on the top of gate dielectric film instead of gate metal of the conventional metal-oxide semiconductor field-effect transistor (MOSFET). Aspects of this work have already been published in a recent conference [13]. However, we conduct wide experiments and extract detailed data results in this new paper. Our 208 & The Institution of Engineering and Technology 2014
biosensors overcome the drift time problem as well, because of the field-enhanced conductivity. Therefore, they will be the most popular low-cost, single-use sensors which can detect not only polar molecules [14–20] but also ions, neutral molecules and biochemical materials, as long as appropriate transducers (recognition layer) are added to the gate dielectric film to provide specificity.
2 Field-effect biosensor fabrication and its sensitive characteristics The biosensor we proposed is a field-effect transistor in which the metallic gate is omitted so that the dielectric material can be exposed to the analyte polymer directly from solution or via a transducer layer with a specific recognition function. The cross-sections of the conventional MOSFET and ChemFET are illustrated comparatively in Figs. 1a and b, respectively. In the conventional MOSFET, the p-type silicon wafer as substrate forms the devices channel, whereas the dielectric film and the aluminium layer form the gate electrode, which are all protected by passivation films. In the ChemFET structure, the aluminium layer which covered the gate dielectric film is not required because the dielectric layer is responsible for the biochemical sensing. Therefore the passivation film is also completely removed in a region delimitated by the area of the source and drain electrodes. To fabricate a ChemFET sensor in Fig. 1b, the process flow is illustrated in Fig. 2, the p-type silicon wafer with a IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
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Fig. 1 Comparative schematic representation of conventional MOSFET and ChemFET a MOSFET with gate metal b ChemFET without gate metal
Fig. 2 Fabrication processes flow of our designed ChemFET sensor (biosensor) IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
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www.ietdl.org thickness of 525 μm is used as a substrate and performs as an active channel in the device operation. First, mask 1# is patterned, by means of conventional lithography and dry etching, to define the p+ guard ring for substrate grounding. Boron (B) is implanted with energy: 40 keV, dosage: 1.0 × 1015 cm−2 and concentration: 1.0 × 1019 cm−3 and the measured sheet resistance is 68 Ω/□. Second, mask 2# is patterned to define the devices drain and source area. Arsenic (As) is implanted with energy: 100 keV, dosage: 4.0 × 1015 cm−2 and concentration: 1.0 × 1019 cm−3 and the measured sheet resistance is 48 Ω/□. Subsequently, mask 3# is patterned to define the devices active area. The 104.3 nm silicon oxide (SiO2) and 18.6 nm silicon nitride (Si3N4) are deposited on the active area to form two layers of complex gate dielectric films. The Si3N4 layer performs a double role in this case: it represents the dielectric film of sensitive gate of biosensors, as well as a stopping layer for the passivation film etching. The thickness of these two layers of gate dielectric film has been chosen to enhance the detection sensitivities. Threshold voltage adjusting is also achieved in this step. Phosphorus (P) is implanted with energy: 150 keV, dosage: 1.5–2.0 × 1013 cm−2 and concentration: 6.5 × 1015 cm−3, thus, the conduction channel occurs at the semiconductor surface under the gate dielectric film. Then, mask 4# is patterned to define the devices Ohmic contact area for drain, source and p-substrate grounding, and mask 5# is patterned to define the metal aluminium connection. Finally, mask 6# is patterned to define the passivation film and open the biosensor’s sensitive window, where the metal and passivation film are all removed from the top of the gate dielectric film. The fabrication process parameters of our design and measurement are listed in Table 1 and a microphotograph of our fabricated biosensor chip is shown in Fig. 3. After cleavage, the die is wire-bonded and encapsulated with a two-component epoxy called EP42HT at a temperature of 300 K or so.
Fig. 3 Microphotograph of our fabricated biosensor chips (portion)
Fig. 4 Schematic diagram showing the bias conditions used for testing the biosensors (ChemFET sensors) with Ag/AgCl reference electrodes
Table 1 Design and measurement processes parameters for our fabricated biosensors No.
Processes
1
boron (B) implantation
2
drain/source implantation
3
drive in
4
phosphorus (P) implantation
5
gate oxidation (SiO2)
6 7
silicon nitride (Si3N4) contact low pressure chemical vapour deposition (LPCVD) (SiO2) passivation LPCVD (Si3N4) atmospheric pressure (AP) CVD (SiO2) APCVD (SiO2) Al/Si Al/Cu
8 9 10 11 12
Controlled parameters 40 keV, 1.0 × 1015 cm−2 100 keV, 4.0 × 1015 cm−2 xj = 3.5 μm 150 keV, 1.5– 2.0 × 1013 cm−2 100 ± 15 nm
Measurement parameters sheet resistance :68 Ω/□ sheet resistance: 48 Ω/□ SiO2 on p+ 59 nm, on n+ 89 nm concentration: 6.5 × 1015 cm−3
20 ± 10 nm 175 ± 17 nm
104.3 nm, Qss 1 × 1010 C/cm2 18.6 nm 189.4 nm
60 ± 6 nm
60 nm
720 ± 72 nm
805.6 nm
175 ± 25 nm 0.4 ± 0.2 μm 1.6 ± 0.2 nm
188.1 nm 0.38 μm 1.5 nm
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Fig. 5 Repeatability and stability testing for our biosensors with detection materials of pure water and S. cercariae protein a Chip 1# b Chip 2# IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
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Fig. 6 Current response related with number of carboxyl group of serial alkyl-acids and different functional groups of serial organic polar acids a Carboxyl group varied for serial alkyl-acids (chip with W/L = 4.20) b Functional group varied for serial organic polar acids (two chips of W/L = 4.20 and W/L = 20.0)
To measure the current response of the detected materials, the terminal bias conditions are illustrated in Fig. 4. The biosensors and an Ag/AgCl reference microelectrode forming a three-terminal device are biased in the same way as a conventional MOSFET in the common source configuration, in which VGS (V) is gate–source bias voltage and VDS (V) is drain–source bias voltage. The output drain– source current, IDS (mA), which indicates the current response for certain detected biomolecules and acids, is recorded by a HP 4156A Precision Semiconductor Parameter Analyzer at a temperature of 300 K. Despite the fact that the conduction mechanisms in biosensors are not clearly elucidated [7, 21–23], their ‘macroscopic’ electronic behaviour may be reasonably described by the same equations as for conventional MOSFETs. At low gate bias voltages, the channel carrier mobility can be assumed as a constant and the current (IDS) flowing through the channel between source and IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
drain is well described by the equation [24, 25]
IDS
2 W VDS = · m · Cins · VGS − Vth VFB · VDS − 2 L W + · s · dsemi · VDS L
(1)
where μ is the channel carriers mobility and σ is the bulk conductivity, Cins is the dielectric capacitance per unit area; dsemi is the thickness of the semiconductor surface depletion layer, VFB is the flat-band voltage and Vth (VFB) is the threshold voltage which is the function of VFB. Therefore the biosensors’ current response of serial polar acids and certain biochemical materials can be attributed to the change in the flat-band voltage induced by the variation of 211
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www.ietdl.org the potential at the electrolyte/gate dielectric interface [26] VFB = Eref + xSOL −
wsemi Qins − − c0 q Cins
(2)
where Eref is the potential of the Ag/AgCl reference electrode, χ SOL is the surface dipole potential of the detected materials, wsemi/q is the semiconductor work function and Qins is the effective charge per unit area and c0 is the voltage drop at the electrolyte/dielectric interface. As shown in (1) and (2), for given devices bias voltage of VDS, VGS and sensitive area parameters of W/L, the devices’ output drain–source current is linearly correlated to χ SOL, c0 (parameters of detected materials) and Qins/Cins, wsemi/q etc. (parameters of fabricated biosensors). In our experiment, biochemical materials combined selected transducers, certain biomolecules and acids can serve as the analyte polymer which is coated on a sensitive area of the biosensor instead of the gate metal by self-assembly techniques (SAM). Thus, work function and charge distribution in gate dielectric film will be changed. Although there is no direct interaction between the semiconductor surface and the analyte polymer, the threshold voltage (Vth) related to semiconductor work function (wsemi/q) is varied. Moreover, because the presence of the analyte polymer has modified the electric field in the surface of the channel, the channel conductivity (σ) is changed accordingly. The change of Vth and σ will vary the biosensor’s output current (IDS) accordingly, which forms the current response of the detected materials. The calculated data of this simple model and the experiments data will be compared in the following section.
Fig. 7 Current response sensitivity related with biosensors sensitive areas (W/L) and current response comparison between simple analytical model calculation and measurement according to variation of haemocyanin concentration a Current response sensitivity related with biosensors sensitive areas (W/L) b Current response comparison between simple analytical model calculation and measurement according to variation of haemocyanin concentration (W/L = 20.0) 212 & The Institution of Engineering and Technology 2014
3 Electrical characterisation results and discussions To optimise the biosensor’s sensitive area, series W/L channel parameters are designed on the same die and fabricated by the same technology. In addition, the devices are made in their depletion operation mode, that is, the devices’ channel is already implanted with an n-type doping agent. Hence, it is always in a weak conducted state and a current will flow through its channel between the drain and the source electrodes even if no gate voltage (VGS = 0.0 V) is applied. Therefore it will show typical resistor output characteristics when VGS = 0.0 V bias conditions are used in Fig. 4. First, we conduct the repeatability and stability test for our fabricated biosensors, by using the measurement schematic diagram shown in Fig. 4, with device parameters of Vth = −1.0 V and device terminal bias conditions of VGS = 0.0 V, VDS = 3.0 V. We selected two biosensor chips with the same sensitive areas of W/L = 20.0 marked as chip 1# and chip 2#, respectively. Then, pure water and protein extracted from Schistosoma cercariae are used as detected materials.
Fig. 8 Current response distribution and comparison among the blank, pure water and S. cercariae worms with two chips a W/L = 4.2 b W/L = 20.0 IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
www.ietdl.org The biosensor’s drain–source response currents for each detected material are measured 10–20 times, in order to verify the repeatability and stability performances of our fabricated biosensor. The process for our measurement is, firstly, coating the detected materials, such as, pure water or protein of S. cercariae on top of the sensor’s gate dielectric film as the analyte polymer with the same SAM formation conditions, and then the sensor’s drain–source current is measured. As shown in Fig. 5, it is clearly demonstrated that the detection experiment is not stable in its initial 1–10 times measurements; however, after this initial period of time, the later detection experiment shows a relatively stable response current on our fabricated biosensor. Then, the detection experiments are conducted for series acids, for example, hexanoic acid, adipic acid and citric acid and so on. We have checked and compared the biosensor’s detection characteristics according to the carboxyl group and functional group of selected acids. Fig. 6a show the measured biosensor’s drain–source current varying with the acids carboxyl group. The blank state measurement data are marked ‘blank’ in the figure. It can be seen that the response current decreases with the increase of carboxyl group for serial alkyl-acid, since there are 1, 2 and 3-carboxyl group for hexanoic, adipic and citric acids, respectively. Those current responses are all smaller than the current response of the blank state. Fig. 6b gives the relationship among detected response current and the functional group of serial polar acid for two sensitive areas of W/L = 4.20 and 20.0. This current response decreases with the order of blank, –OOH, –OH and –NH2 for both sensitive areas of W/L = 4.20 and 20.0. The biosensor’s quantified sensitivity [27] is measured according to the series sensitive area and different
concentrations of target detected biomolecules and acids. Haemocyanin, the protein extracted from snails, is used in this experiment. Fig. 7a shows the current response of haemocyanin for series W/L values, here the sensitivity is defined as response current per W/L values, that is, sensitivity = IDS/(W/L). It is clear that the sensitivity decreases with the increase of W/L. Fig. 7b demonstrates the current response with different haemocyanin concentrations from 0.1753 to 0.7109 (mg/ml) for sensitive area of W/L = 20.0 chips. The measurement data are also compared with the calculated data from the simple model represented by formula (1), in which the main model parameters used in our calculation are listed as W/L = 20, VDS = 3 (V), μn = 450 (cm2/V/s), dsemi = 30 (nm), gate oxide film (SiO2): tox = 104.3 (nm), εr = 3.9, gate silicon nitride film (Si3N4): tox = 18.6 (nm), εr = 6.2, Nsub = 1.7 × 1016 (cm−3), Vth = −1.0 (V) and σ = q·nsub·μn (Siemens/cm). We assume that, for each measured target molecule concentration, Vth and σ will vary and its value will be determined by measurement and calculation. Fig. 7b shows that the higher the haemocyanin concentration, the larger response current we can detect, and those response currents are all larger than that of blank. The next detection experiment is conducted for S. cercariae worms [28] with different sensitive area, the measurement results are compared with that of pure water and blank, as shown in Fig. 8, where (a) is for W/L = 4.20 and (b) is for W/L = 20.0. It can be seen that the response current for pure water and blank is relatively stable under certain experimental conditions. However, the response currents for S. cercariae worms are distributed in a certain current range. The reason for this data distribution range is (i) the variety of density for S. cercariae worms, (ii) the
Fig. 9 Current response for detection of S. cercariae worms using linoleic acids as transducers (recognition) layer a W/L = 4.20 b W/L = 20.0 IET Nanobiotechnol., 2014, Vol. 8, Iss. 4, pp. 208–215 doi: 10.1049/iet-nbt.2013.0061
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www.ietdl.org different positions on top of the sensitive area for each S. cercariae worms after finishing coating on the dielectric film, (iii) the characteristics of SAM S. cercariae worms and its formation conditions is related to each SAM processing. Since there are no or little electric dipole characteristics in pure water, the response current of pure water is positive and close to that of blank. However, S. cercariae worms have high dipole charge characteristics, properties of unique electronic conduction and delocalised electronic structure along the conjugated backbone, and its measured response currents are negative and far apart from that of pure water and blank. In all the above experiments, the effect of surface charging state density and especially pH value of pure water to the biosensor’s response current must be considered in our experiment for blank and pure water test. We assume that the influence of surface charging state density is already consistent in the output current response. However, water’s pH value has large influence on the surface charging state density, the surface charging state density will increase with the increase of pH value [29]. Hence, we must ensure the pH value to be stable during our experiment, that is, using the same barrel/batch of pure water with pH value of 6.8 ± 0.2. Otherwise, our detection results comparison will be invalid since the pure water pH value is varied with temperature and environment. For example, the research results in [29] show that the surface charging state density varies rapidly when the pH value varies from 5 to 7, which will make the measurement data unstable and the experiment uncontrollable. The research results in [29] also show that, the surface charging state density changes slowly when the pH value is smaller than 5. Hence, serial polar acids can be used as transducers [1, 30] to detect S. cercariae worms. For example, we use linoleic acid as transducer materials in one of our experiments. Fig. 9 shows our experiment in repeated detections of S. cercariae worms by using linoleic acid as a transducer film. The test processes are repeatedly carried out seven times. For each time of measurement, we first make linoleic acid transducers using SAM processes, and measure the sensor response current for this transducer. Then, the S. cercariae worms are coated on top of the pervious transducer film also by SAM processes. The synthesis-induced response current is measured and compared with the current induced only by the transducer film. As shown in Fig. 9, (a) is for W/L = 4.20 and (b) for W/L = 20.0. Generally, after the SAM S. cercariae worms, the synthesis-induced response currents are larger than that of linoleic acid transducers except for the second time measurement in Fig. 9b.
4
Conclusions
We have designed and fabricated ChemFET sensors (biosensors) on p-type silicon substrates with channel parameters of W/L = 4.2, 20.0 and so on. We have also investigated the ability of our fabricated biosensors to perform charge detection in aqueous media. We conduct the repeatable and stable experiments 10–20 times, using pure water and S. cercariae protein as detected materials and compared their response current results. The detecting ability for series acids is verified by comparing the sensor’s current response for specifically selected acid with different carboxyl group and functional group. The detected results are also compared with that of blank. The biosensor’s sensitivity is verified by checking the variation of the 214 & The Institution of Engineering and Technology 2014
current response with serial sensitive area (W/L) and serial haemocyanin concentrations, which is compared with simple model calculation data. After finishing the above experiment, we selected two groups of sensors for S. cercariae worms’ detection, and the detected response currents are compared with that of pure water and blank. For sensitive areas of W/L = 4.2 and 20.0, the current range for S. cercariae worms are 0.145–0.060 mA and 1.650– 1.300 mA, respectively. The minimum current differences (resolution spacing) compared with blank were 0.075 and 0.350 mA, respectively. The transducer (recognition) layer experiment is also conducted using linoleic acid as transducer for detecting S. cercariae worms, which gives clearly detected resolution in our experiment.
5
Acknowledgments
The authors would like to thank Prof. and PhD J.Q. Han and senior engineer Mr. X.L. Wang in the Microelectronics Institute for their support of wafer fabrication and valuable feedback. This research is supported by the Natural Science Fund Committee of China (NSFC) (no. 51377025, GJJ13661).
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