Low-Cost Nanoribbon Sensors for Protein Analysis in Human Serum

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Low Cost Nanoribbon Sensors for Protein Analysis in Human Serum Using a Miniature Bead-Based Enzyme-Linked Immunosorbent (ELISA) Assay Chunxiao Hu, Ioannis Zeimpekis, Kai Sun, Sally Anderson, Peter Ashburn, and Hywel Morgan Anal. Chem., Just Accepted Manuscript • DOI: 10.1021/acs.analchem.6b00702 • Publication Date (Web): 01 Apr 2016 Downloaded from http://pubs.acs.org on April 3, 2016

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Analytical Chemistry

Low Cost Nanoribbon Sensors for Protein Analysis in Human Serum Using a Miniature Bead-Based Enzyme-Linked Immunosorbent (ELISA) Assay Chunxiao Hu,† Ioannis Zeimpekis,† Kai Sun†, Sally Anderson,‡ Peter Ashburn,† and Hywel Morgan*,† †

Department of Electronics and Computer Science, and Institute for Life Sciences, University of Southampton, United Kingdom

‡

Sharp Laboratories of Europe, Oxford, United Kingdom

ABSTRACT: We describe a low cost thin-film transistor (TFT) nanoribbon sensor for detection of the inflammatory biomarker Creactive protein (CRP) in human serum via a miniature bead-based Enzyme-Linked Immunosorbent Assay (ELISA). The TFT nanoribbon sensor measures the reaction products from the ELISA via pH changes. The bead-based ELISA decouples the protein functionalization steps from the sensor surface, increasing the signal and simplifying the assay. The ability to directly sense proteins in human serum in this way overcomes the Debye length limitation associated with nanowire and nanoribbon biosensors. Compared to classically fabricated nanowires, the TFT nanoribbon sensors are simple, extremely easy to fabricate and should therefore be much cheaper to manufacture. TFT nanoribbon sensors, configured to measure pH were used for quantitative detection of CRP spiked into human serum at concentrations as low as 0.2 ng/mL, which is ten thousand times lower than needed for diagnostic purposes, providing the potential for applications that require very high sensitivity.

Quantification and analysis of biological processes, especially determination of low concentrations of proteins, are important in healthcare.1 The most widely used assay is the Enzyme Linked Immunosorbent Assay or ELISA2 which uses an enzyme to amplify the biochemical binding event. Surface Plasmon Resonance (SPR)3 sensors on the other hand are label free but usually require expensive optical equipment for detection. Recently, Field Effect Transistor (FET) devices such as silicon nanowire and nanoribbon biosensors have been developed for direct, high sensitivity, label-free sensing of biomolecules. These devices are small and could be integrated into point of care (PoC) systems.4-8 Nanowire/nanoribbon devices configured as ion-sensitive FETs (ISFETs) have been used for a large variety of applications, including pH sensing,4,9-12 DNA sensing,13-19 protein sensing,4,7,8,16,20-25 and enzyme detection.26 However, the technology has not yet evolved into a robust platform for analysis of samples in human body fluids such as blood due to several limitations. One is the limitation on suspending medium composition imposed by the screening effect of the electrical double layer. Label free FET mediated biosensing of biomolecules is based on the fact that the charge of a molecule bound to the surface of the transistor gate changes the surface potential, thus varying the carrier density inside the channel. In physiological media, the Debye length is around 0.7 nm; any charge from a bound molecule beyond this distance will be effectively screened by the medium. Several methods have been reported to overcome this limitation. Generally the approach is to perform measurements in a low salt concentration (e.g. 1 mM) buffer, thereby increasing the Debye screening length.8,27,28 Other methods include the use of short molecular receptors to reduce the distance between the surface and the analyte,29,30 or

modification of the sensor surface with PEG polymers to increase the effective screening length.31 Recently, these devices have been used to measure proteins through a modified ELISA where the electronic readout is via pH change.32,33 However, as for the label-free assays, the receptor molecules were linked directly to the sensor surface, for example immobilized to the gold leads near the nanowire gate or directly onto the nanoribbon itself. These approaches limit the flexibility of the sensor and also the sensitivity and modify the kinetics. It also complicates manufacturing since each sensor has to be specifically functionalized with a different capture moiety. Here we report a more flexible approach for the detection of C-reactive protein (CRP) in human serum using a miniature bead-based ELISA with pH readout coupled with a low cost TFT nanoribbon sensor for readout. Recently a magnetic beads based assay was demonstrated in a microfluidic chip with in-situ washing. Negatively charged single stranded DNA was used as a label to modify the charge density and the signal read out when the beads were attracted to the surface of a metal oxide using a magnet.34 However, the use of miniature beads coupled with enzyme amplified pH readout provides a highly flexible assay format that is simple and uses disposable electronics to directly read out the signal. The use of beads as carriers also makes the system scalable. For example large area very low cost TFT arrays containing hundreds or thousands of sensors could be used to measure multiple analytes simultaneously. CRP is one of the most important acute-phase proteins that exists in blood and is synthesized by the liver.35 It is widely used as a general biomarker for inflammation and infection, and has become a good indicator for evaluating risks of cardiovascular diseases.36,37 Generally,

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the CRP concentration in blood is less than 10 mg/L for healthy individuals, but can increase 1000 fold during an acute phase of inflammation.38 Fast detection of CRP has been demonstrated using FET biosensors with pH readout,39 but the proteins were immobilized onto a gold sensor surface and the lowest concentration of CRP that could be detected was 29 ng/ml. In this paper we have modified a conventional beadbased ELISA to produce a pH change that is read using the TFT nanoribbon sensor. Since molecular detection is measured indirectly through an enzyme-substrate reaction, the Debye screening effect is eliminated and both specificity and sensitivity are increased. Furthermore, the assays are implemented using magnetic beads on the substrate rather than the sensor surface, which increases the capture cross-section and improves the reaction rate. It also provides significant advantages in terms of manufacturing biosensors since the sensors do not need to be chemically functionalized – the chemistry is entirely confined to the carrier beads, which can be introduced into the sensor prior to the assay. The ureaurease reaction was first used as a model system to evaluate the sensitivity of the TFT nanoribbon sensors for enzymesubstrate detection. The same enzyme was then used to detect CRP in human serum using a miniature bead-based ELISA. The urease-induced pH change was monitored to determine the quantity of protein in the serum. In fact, any bead-based ELISA could be read out in this way. Compared to conventional nanowire fabrication techniques, e.g. topdown15,40,41 or bottom–up4,42,43 approaches, the polysilicon TFT nanoribbon sensor was fabricated with a simple three-mask approach,44 which produces a technology that is simple, cheap and suitable for the mass manufacture of disposable biosensors for PoC applications.

EXPERIMENTAL Device fabrication and measurement configuration. A three-mask fabrication method44 was used to manufacture the TFT nanoribbon sensors shown in Fig. 1. Briefly the fabrication process involved defining the polysilicon transistor using photolithography, contacting Source and Drain using metal tracks followed by deposition and patterning of the thick resist SU8 to create a sensing window. Fabrication details have been described previously.44 The TFT nanoribbon sensor has two gates. The top gate is a liquid gate controlled by applying a potential to an integrated on-chip Ag/AgCl electrode (Fig. 1a). The bottom gate is the back of the substrate, which is grounded for all the measurements described in this paper. Silicon dioxide is used as the gate dielectric, which is in contact with the liquid. An SU8 layer insulates all the contacts and defines the sensing windows (Fig. 1b) where the TFT nanoribbon sensors are in contact with test solution, which is confined with a small well made from PDMS. A dual-channel picoammeter/voltage source (Model 6482, Keithley) and a connection jig (Sharp Laboratories of Europe) were used to apply voltage and acquire data.

Figure 1. (a) Schematic diagram showing the small droplet of Microscope image of one sensor, µm long) connected in parallel electrode.

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of the TFT nanoribbon sensor sample (not to scale); (b) containing 30 nanoribbons (40 with the integrated reference

Enzyme reaction measurement. A simple enzyme-substrate reaction was demonstrated using urea (Sigma) and urease (Sigma). Five different concentrations of urea were prepared in a measurement buffer (0.01×PBS + 150 mM NaCl).32 Prior to each measurement, the entire sensing area (including the silver/silver chloride reference electrode) was incubated in 3% bovine serum albumin (BSA) in PBS to minimize non-specific binding. Urea solution (100 µL at various concentrations) was pipetted into the sensing window and the source-drain current monitored against time (source-drain voltage Vds = 0.1 V; Liquid gate voltage Vlg = 0.05 V) to provide a baseline signal. Subsequently urease (50 µL at 0.45 mg/mL) was added and the solution was quickly mixed by pipetting up and down several times. Urease catalyzes the hydrolysis of urea to ammonia and carbon dioxide leading to an increase in the pH as the reaction proceeds. This is read as a decrease in the source-drain current (for n-type TFT nanoribbon sensors). Bead functionalization and measurement. A bead based electronic ELISA was implement using magnetic beads that were modified sequentially as shown in Fig. 2a. Two different assays were performed. In one, the entire chemistry was performed in a test tube “off-chip” and the products of the reaction measured with the TFT nanoribbon sensor. In the other, the entire reaction was performed in a 3µL volume “onchip” Commercial superparamagnetic beads (3 µm in diameter) with CRP capture antibody were purchased from R&D systems. These beads were re-suspended in PBS to a concentration of 5000 beads/µL, which provides sufficient binding sites for the protein. For the off-chip assay, an aliquot (10 µL) of these beads was mixed with a solution of CRP (R&D systems) (10 µL) in the well of a microplate. The CRP stock was prepared by spiking protein into human serum (Sigma) to give a concentration of 10 mg/L CRP. This is the normal concentration in the blood of a healthy person. This was diluted with PBS to produce the desired concentrations for the assay. The bead suspension was incubated at room temperature for 30 minutes on a horizontal orbital microplate shaker set at 800±50rpm. The beads were then washed three times in PBS by concentrating the beads with a magnet and aspirating the supernatant. Biotinylated CRP detection antibody solution (10 µL, 0.1 mg/ml in PBS; R&D systems) was added to the well and incubated at room temperature for 10 minutes with shaking, followed by three washes. Streptavidin (10 µL, 0.1 mg/ml in PBS; R&D systems) was then added to the well and incubated at room temperature for

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another 10 minutes on the shaker, followed by three washes. Finally, biotinylated urease (10 µL, 0.2 mg/mL; Insight Biotechnology) was added and the solution incubated at room temperature for 10 minutes followed by further bead washing. The beads were then re-suspended in the working buffer (0.01×PBS + 150 mM NaCl) to a volume of 3 µL and transferred to the sensor chip.

at room temperature for 2 hours on a horizontal orbital microplate shaker. Beads were washed with PBS before adding biotinylated CRP detection antibody solution (100 µL, 0.1 mg/ml in PBS), followed by incubation and washing. Protein concentration was assayed by adding StreptavidinHRP (100 µL, 0.1 mg/ml in PBS; Thermo Fisher Scientific), followed by incubation followed by three washes. TMB substrate solution (100 µL; Thermo Fisher Scientific) was then added and incubated for 30 minutes, followed by a stop solution (50 µL; Thermo Fisher Scientific). Since the magnetic beads could affect the optical density of the fluid, the solution was transferred to another well leaving the beads in place. The optical density (O.D.) of each well was measured by a microtiter plate reader with a filter set to 450 nm.

RESULTS AND DISCUSSION

Figure 2. Schematic diagram of the magnetic bead sandwich assay. (a) Magnetic beads with CRP capture antibody were sequentially incubated with CRP at different concentrations, followed by biotinylated CRP detection antibody, streptavidin, and biotinylated urease. (b) Urea was added to initiate the reaction. The change of pH is related to the concentration of CRP and was detected by the TFT nanoribbon sensor.

After adding the magnetic bead suspension, the source-drain current signal (source-drain voltage Vds = 0.1 V; liquid gate voltage Vlg = 0.05 V) was measured as a function of time. Then urea solution (3 µL, 10 µM diluted with working buffer) was added and rapidly mixed, and the pH change measured from the Source-Drain current. The same assay was also carried out with the entire reaction implemented in the sensor well “on-chip”. In this case, beads were suspended in PBS to a concentration of 5000 beads/µL and a 3-µL aliquot added to the sensor well. A solution of CRP in human serum (3 µL) was added by pipetting and the mixture mixed by successive pipetting and incubated at room temperature for 30 minutes. The beads were then washed three times in PBS by concentrating the beads with magnet and aspirating the supernatant. Biotinylated CRP detection antibody solution (3 µL, 0.3 mg/ml in PBS) was added to the well, mixed and incubated at room temperature for 10 minutes, followed by three washes. Streptavidin (3 µL, 0.3 mg/ml in PBS) was then added to the well, mixed and incubated at room temperature for another 10 minutes, followed by three washes. Finally, biotinylated urease (3 µL, 0.6 mg/mL) was added, mixed and incubated at room temperature for 10 minutes followed by washing. The bead suspension was then re-suspended by adding the working buffer to a volume of 3 µL. A conventional colorimetric bead based ELISA was also carried out. In this case beads were suspended in PBS to a concentration of 5000 beads/µL. An aliquot (100 µL) of these beads was mixed with a 100 µL solution of CRP in human serum in the well of a microplate. The mixture was incubated

Electrical characteristics. The pH sensitivity of TFT nanoribbon sensors with a silicon dioxide dielectric surface was first characterized from IdsVlg sweeps with calibrated pH buffers. As shown in Figure 3a, four different pH buffers (pH = 9.0, 7.0, 5.0 and 3.0 respectively) were used to characterize the electrical response of the devices by sweeping the liquid gate voltage and recording the source-drain current (Vds = 0.1 V). The TFT nanoribbon sensors are n-type and therefore the source-drain current Ids increased with decreasing pH. The transconductance gm = ∂Ids/∂Vlg is used to characterize individual TFT nanoribbon sensors, and this value was extracted from the linear region of the IdsVlg curve (-0.1 to 0.1 V in this case) and found to be nearly constant at ~2×10-7 A/V for all four pH values. The threshold voltage shift ∆Vth = Ids/gm calculated from the IdsVlg curves is shown in Fig. 3b. The figure shows that for liquid gate voltages ranging from -0.1 to 0.1 V, ∆Vth is a linear function of pH with a slope of ~ 33 mV/pH in line with literature values for SiO2 surfaces.45

3



becomes more basic. The reaction rate was determined from the initial slope of the reaction (dashed square in Fig. 4a) and fitted to the Michaelis–Menten equation (Fig. 4b):

(a) 10

-6

[]

10-7

is the reaction rate (current or threshold voltage change with time), Vmax is the maximum rate and KM, is the Michaelis constant.

pH9 pH7 pH5 pH3

Vds = 0.1 V gm = 2.01×10-7 A/V

10

10-8 0

1

2

0

Vlg (V)

-10

1.6 -100 mV -60 mV -20 mV 0 mV 20 mV 60 mV 100 mV

1.5

Vlg = 1.4

(Id-Id0)/gm (mV)

(b)

(2)

[]

0 mM Urea 2 mM Urea 5 mM Urea 10 mM Urea 20 mM Urea 30 mM Urea

-20 -30

pH

Ids (A)

=

∆Vth (V)

-40 -50

1.3

1.2

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-60 -70

Vds = 0.1 V Slope = 33.5±1.9 mV/pH 3

4

5

6

0

200

400

600

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1000

Time (s) 7

8

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pH

Figure 3. (a) TFT Nanoribbon sensor source-drain current (Ids) vs liquid gate voltage measured for different pH buffer solutions. The transconductance gm=∂Ids/∂Vlg is extracted from the linear region. (b) Threshold voltage change ∆Vth for 7 different liquid gate voltages. The voltage is a linear function of pH with a subNernstian slope of 33 mV/pH.

Enzyme reactions. A simple enzyme-substrate reaction was demonstrated using urea-urease. Urea serves an important role in metabolic processes and is the main nitrogen-containing substance in the urine of mammals. Its concentration is an important indicator of some diseases for example heart and renal failure.46,47 Urease catalyzes the hydrolysis of urea into one carbon dioxide and two ammonia molecules increasing the pH: ( ) + → + 2 (1) Urease (at 0.45 mg/mL) was pipetted into a well around the TFT nanoribbon sensors, and the change in pH was measured upon the addition of urea at six different concentrations (0 to 30 mM). Figure 4a shows the recorded electrical readout when a small volume of urea solution (100 µL) was added to the sensing window (point 1 on graph). As shown, the signal remains stable with time. After 300 s urease (50µL) was added and mixed (point 2 on graph) and the electrical signal monitored (for typically 10 minutes). The transconductance gm obtained during the electrical characterization of a single device was used to normalize the response from different devices. The data shows that the source-drain current decreases as the enzyme reaction proceeds and the solution

Reaction Rate (µ µ M/min)

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4

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pH meter Nanoribbon

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0 0

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Concentration (mM)

Figure 4. Threshold voltage vs time for the urea-urease reaction measured with an n-type TFT nanoribbon sensor. (a) Source-drain current normalized against transconductance (left y-axis) for different urea concentrations with device operating in subthreshold. Number 1 indicates the point at which urea was added to the sensor, and 2 when urease was added. Calculated equivalent pH changes are shown on the right y-axis. (b) Michaelis–Menten curves obtained from change in current during the first few minutes (dashed square in Figure 4a). The same reaction was also carried out in a test tube and the rate measured using a pH meter. Data are the mean + SEM (three measurements with two devices).

For comparison, the urea/urease reaction was also monitored in a larger volume using a standard pH meter. Data were recorded every 5 seconds for the first 5 minutes and every 30 seconds for the following 10 minutes. The Michaelis constants (KM) were calculated and found to be 15.9 ± 2.1 mM (TFT nanoribbon sensor) and 9.2 ± 0.7 mM (pH meter) respectively. This can be compared with the literature value of 25 mM48 (Table 1). The KM determined using a TFT nanoribbon sensor is slightly higher than measured conventionally, this may be

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due to differences in mixing or temperature between the two systems. The turnover number (kcat) was also determined from =

[ ]!

(3)

where [E] 0 is the enzyme concentration. Turnover numbers were 3.7×104 s-1 and 3.4×104 s-1 for the TFT nanoribbon sensor and pH meter respectively, which compares with literature values of around 104 s-1.49,50 Table 1 summarizes the data. These measurements indicate that the TFT nanoribbon sensors can be used to reliably measure enzyme-substrate reaction rates. Bead-based ELISA. In a bead-based ELISA, the capture antibodies are immobilized on the surface of small beads, rather than the plastic surfaces of microtiter plates. Using beads rather than a planar surface significantly increases the surface area available for capture, and also makes mixing easier thereby speeding up the reaction. These assays require relatively small sample volumes and demonstrate increased sensitivity.51 For this assay, magnetic beads functionalized with CRP capture antibodies were used as capture surfaces. The assay was read out by adding urea to the bead suspension, leading to an increase in the pH of the solution that is proportional to the amount of captured CRP. Seven concentrations of CRP spiked into human serum were measured, ranging from 0 to 500 ng/ml. The same reaction was performed in two different ways. Either “off-chip” in a small tube or entirely “on chip”. In this case the beads were washed on chip by concentrating them with a small hand held magnet (see experimental section). The enzyme reaction was measured through the change in current as a function of time and the data are shown in Figure 5. The source-drain current

was continuously measured and the bead suspension added at point 1. Immediately after adding urea to the bead suspension at point 2, a sharp decrease in voltage was observed (Fig. 5a) caused by agitation of the liquid when the suspension was mixed. The signal stabilized and the voltage decreased after around 30 seconds. This time point was chosen as the beginning of the reaction. The change in threshold voltage is plotted in the figure and the slope of the curve over the first minute (marked by a red dashed square) was used to calculate the reaction rate (using Matlab). A plot of initial reaction rate against CRP concentration is shown in Figure 5b. The end point voltage after 500 seconds was also measured. In total, eight sets of measurements were performed on the TFT nanoribbon sensors, four with the entire assay performed on-chip (red curve) and four with the end product of the reaction measured on chip (black curve). The data was fitted to the Hill equation (Fig. 5b) to obtain the equilibrium binding constant Kd which describes the affinity of the ligand for the protein. In this case it is assumed that the binding constant is dominated by the interaction between the CRP and the antibody. The Kd was determined for each separate experiment and the mean binding constant determined as 43 ± 10 ng/ml for off-chip and 38 ± 13 ng/ml for on-chip. Figure 5c shows a similar standard curve extracted from the end point voltage change instead of the initial rates. This gave slightly lower binding constants of 30 ± 11 ng/ml for off-chip and 29 ± 10 ng/ml for on-chip assay. To verify these binding constants, a conventional colorimetric end-point bead based ELISA was performed (Fig 5d), (n = 3). This gave a binding constant Kd = 43± 6 ng/ml, which is similar to that measured on chip.

Table 1. Comparison of KM and kcat values for the urea-urease reaction determined with the TFT nanoribbon sensors (three measurements with two devices) and also a pH meter in bulk.

KM (mM)

TFT Nanoribbon Sensor

pH Meter

Literature

15.88 ± 2.09

9.18 ± 0.69

25 48

R2

0.9937

0.9984

-

kcat (s-1)

3.724×104

3.356×104

3×104 49

5



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Figure 5. Detection of C-reactive protein using bead-based ELISA on TFT nanoribbon sensors. (a) Change in threshold voltage for different concentrations of CRP (with on-chip chemistry). The change in source-drain current is measured and converted to potential shift by dividing by the transconductance gm. The equivalent change in pH is shown on the right y-axis. Number 1 indicates the time point when the beads are pipetted onto the sensor surface, and point 2 when urea was added. (b) Plot of the initial reaction rate (determined within the dashed square in Figure 5a) vs concentration of protein. Red curve is data for on-chip assay; black curve for off-chip assay (c) Plot of the final voltage (after 500 sec) against protein concentration. The plots were fitted to the Hill equation. Data are the mean ± SD (four measurements with two devices). (d) Standard curve measured from the endpoint of a conventional colorimetric bead based ELISA. Optical density measured with a plate reader.

The results from on-chip or off-chip chemistry are very similar and indicate that the concentration of CRP in human serum can be measured at a concentration as low as 0.2 ng/ml in a volume of only 3 µL. The Limit of Detection (LOD) determined from 3 x standard deviation of the blank signal is approximately 0.05 ng/ml. The total assay time for the on-chip chemistry protocol is approximately 1 hour, which could be reduced further with optimization of the fluidics and development of magnetic bead agitation protocols for in-situ continuous mixing. The method presented here is applicable to the detection of many other proteins in high salinity buffers such as human serum (where optimized bead based ELISAs are available). Scaling up the system could provide a simple but cost-effective approach for analyzing large numbers (thousands) of different proteins using arrays of thin film transistors.

CONCLUSION AND OUTLOOK Low cost thin-film transistor nanoribbon sensors with integrated reference electrodes have been fabricated and used to analyze enzyme-substrate reactions via pH changes. The device was tested using the urease-urea reaction to characterise enzyme properties. An assay for the quantitative detection of the inflammatory biomarker C-reactive protein in human serum has been developed using a miniature beadbased ELISA with pH readout. The protein assay can be performed using the TFT nanoribbon sensors in high ionic strength buffer, unlike conventional label-free approaches that require low-ionic strength and a large Debye length. The ureaurease system demonstrates that enzyme kinetics can be reliably analyzed using the sensor. A magnetic bead-based ELISA was used to detect CRP in human serum at concentrations down to 0.2 ng/mL in a volume of 3 µL. The

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entire assay, including functionalization, mixing and detection was performed on the sensor in under an hour. Compared with current analytical methods for CRP quantification, no expensive detection equipment is required and the volume of the test sample is very small. An important advantage of this assay is that all the functionalization steps are performed on the magnetic beads, not on the sensor surface. This simplifies manufacturing and assay development and also allows the device to be reused. Many different bead-based ELISA are commercially available, therefore minimal surface chemistry development or assay optimization is required in transferring these to the device Further optimization should lead to smaller assay volumes, speeding up the reaction and enabling the detection of very low amounts of proteins in small volumes of serum. This electronic ELISA brings low-cost TFT nanoribbon sensor technology a step closer to a point of care diagnostic system.

AUTHOR INFORMATION Corresponding Author * Email: [email protected]

ACKNOWLEDGMENT The authors would like to acknowledge the Technology Strategy Board (TSB) and the Engineering and Physical Sciences Research Council (EPSRC: EP/K502327/1) for funding this work. We would also like to thank Gregory Gay, Ben Hadwen, Chris J. Brown and Jonathan Buse of Sharp Laboratories Europe for many useful discussions and development of a measurement jig

REFERENCES (1) Chen, K. I.; Li, B. R.; Chen, Y. T. Nano Today 2011, 6, 131-154. (2) Engvall, E.; Perlmann, P. Immunochemistry 1971, 8, 871-874. (3) Zeng, S.; Baillargeat, D.; Ho, H. P.; Yong, K. T. Chem Soc Rev 2014, 43, 3426-3452. (4) Cui, Y.; Wei, Q.; Park, H.; Lieber, C. M. Science 2001, 293, 1289-1292. (5) Tarasov, A.; Wipf, M.; Stoop, R. L.; Bedner, K.; Fu, W. Y.; Guzenko, V. A.; Knopfmacher, O.; Calame, M.; Schonenberger, C. Acs Nano 2012, 6, 9291-9298. (6) Rajan, N. K.; Duan, X. X.; Reed, M. A. Wires Nanomed Nanobi 2013, 5, 629-645. (7) Elfstrom, N.; Karlstrom, A. E.; Linnros, J. Nano Lett 2008, 8, 945949. (8) Hakim, M. M. A.; Lombardini, M.; Sun, K.; Giustiniano, F.; Roach, P. L.; Davies, D. E.; Howarth, P. H.; de Planque, M. R. R.; Morgan, H.; Ashburn, P. Nano Lett 2012, 12, 1868-1872. (9) Knopfmacher, O.; Tarasov, A.; Fu, W. Y.; Wipf, M.; Niesen, B.; Calame, M.; Schonenberger, C. Nano Lett 2010, 10, 2268-2274. (10) Chen, S. Y.; Bomer, J. G.; Carlen, E. T.; van den Berg, A. Nano Lett 2011, 11, 2334-2341. (11) Peretz-Soroka, H.; Pevzner, A.; Davidi, G.; Naddaka, V.; Tirosh, R.; Flaxer, E.; Patolsky, F. Nano Lett 2013, 13, 3157-3168. (12) Zeimpekis, I.; Sun, K.; Hu, C. X.; Thomas, O.; de Planque, M. R. R.; Chong, H. M. H.; Morgan, H.; Ashburn, P. Nanoscale Res Lett 2015, 10. (13) Gao, A. R.; Lu, N.; Wang, Y. C.; Dai, P. F.; Li, T.; Gao, X. L.; Wang, Y. L.; Fan, C. H. Nano Lett 2012, 12, 5262-5268. (14) Hahm, J.; Lieber, C. M. Nano Lett 2004, 4, 51-54. (15) Li, Z.; Chen, Y.; Li, X.; Kamins, T. I.; Nauka, K.; Williams, R. S. Nano Lett 2004, 4, 245-247. (16) Bunimovich, Y. L.; Shin, Y. S.; Yeo, W. S.; Amori, M.; Kwong, G.; Heath, J. R. J Am Chem Soc 2006, 128, 16323-16331.

(17) Zhang, G. J.; Zhang, G.; Chua, J. H.; Chee, R. E.; Wong, E. H.; Agarwal, A.; Buddharaju, K. D.; Singh, N.; Gao, Z. Q.; Balasubramanian, N. Nano Lett 2008, 8, 1066-1070. (18) Xie, P.; Xiong, Q.; Fang, Y.; Qing, Q.; Lieber, C. M. Nat Nanotechnol 2012, 7, 119-125. (19) Hao, Y. Q.; Zhou, B. B.; Wang, F. B.; Li, J.; Deng, L.; Liu, Y. N. Biosensors & Bioelectronics 2014, 52, 422-426. (20) Zheng, G. F.; Lieber, C. M. Methods Mol Biol 2011, 790, 223237. (21) Stern, E.; Klemic, J. F.; Routenberg, D. A.; Wyrembak, P. N.; Turner-Evans, D. B.; Hamilton, A. D.; LaVan, D. A.; Fahmy, T. M.; Reed, M. A. Nature 2007, 445, 519-522. (22) Kumar, M. A.; Jung, S.; Ji, T. Sensors-Basel 2011, 11, 50875111. (23) Yu, R. M.; Pan, C. F.; Wang, Z. L. Energ Environ Sci 2013, 6, 494-499. (24) YacoubGeorge, E.; Wolf, H.; Koch, S.; Woias, P. Sensor Actuat B-Chem 1996, 34, 429-434. (25) Selvanayagam, Z. E.; Neuzil, P.; Gopalakrishnakone, P.; Sridhar, U.; Singh, M.; Ho, L. C. Biosens Bioelectron 2002, 17, 821-826. (26) Mu, L. Y.; Droujinine, I. A.; Rajan, N. K.; Sawtelle, S. D.; Reed, M. A. Nano Lett 2014, 14, 5315-5322. (27) Stern, E.; Wagner, R.; Sigworth, F. J.; Breaker, R.; Fahmy, T. M.; Reed, M. A. Nano Lett 2007, 7, 3405-3409. (28) Duan, X. X.; Li, Y.; Rajan, N. K.; Routenberg, D. A.; Modis, Y.; Reed, M. A. Nat Nanotechnol 2012, 7, 401-407. (29) Elnathan, R.; Kwiat, M.; Pevzner, A.; Engel, Y.; Burstein, L.; Khatchtourints, A.; Lichtenstein, A.; Kantaev, R.; Patolsky, F. Nano Lett 2012, 12, 5245-5254. (30) Maehashi, K.; Katsura, T.; Kerman, K.; Takamura, Y.; Matsumoto, K.; Tamiya, E. Anal Chem 2007, 79, 782-787. (31) Gao, N.; Zhou, W.; Jiang, X. C.; Hong, G. S.; Fu, T. M.; Lieber, C. M. Nano Lett 2015, 15, 2143-2148. (32) Stern, E.; Vacic, A.; Li, C.; Ishikawa, F. N.; Zhou, C. W.; Reed, M. A.; Fahmy, T. M. Small 2010, 6, 232-238. (33) Aroonyadet, N.; Wang, X. L.; Song, Y.; Chen, H. T.; Cote, R. J.; Thompson, M. E.; Datar, R. H.; Zhou, C. W. Nano Lett 2015, 15, 1943-1951. (34) Lin, Y. H.; Peng, P. Y. Analytica Chimica Acta 2015, 869, 3442. (35) Thompson, D.; Pepys, M. B.; Wood, S. P. Struct Fold Des 1999, 7, 169-177. (36) Miller, V. M.; Redfield, M. M.; McConnell, J. P. Curr Vasc Pharmacol 2007, 5, 15-25. (37) Mygind, N. D.; Harutyunyan, M. J.; Mathiasen, A. B.; Ripa, R. S.; Thune, J. J.; Gotze, J. P.; Johansen, J. S.; Kastrup, J.; Grp, C. T. Inflamm Res 2011, 60, 281-287. (38) Clyne, B.; Olshaker, J. S. J Emerg Med 1999, 17, 1019-1025. (39) Lin, C. L.; Kang, Y. W.; Chang, K. W.; Chang, W. H.; Wang, Y. L.; Lee, G. B. 2014 Ieee 27th International Conference on Micro Electro Mechanical Systems (Mems) 2014, 250-253. (40) Lin, H. C.; Lee, M. H.; Su, C. J.; Huang, T. Y.; Lee, C. C.; Yang, Y. S. Ieee Electr Device L 2005, 26, 643-645. (41) Wenga, G.; Jacques, E.; Salaun, A. C.; Rogel, R.; Pichon, L.; Geneste, F. Biosensors & Bioelectronics 2013, 40, 141-146. (42) Whang, D.; Jin, S.; Wu, Y.; Lieber, C. M. Nano Lett 2003, 3, 1255-1259. (43) Zheng, G. F.; Patolsky, F.; Cui, Y.; Wang, W. U.; Lieber, C. M. Nature biotechnology 2005, 23, 1294-1301. (44) Sun, K.; Zeimpekis, I.; Lombardini, M.; Ditshego, N. M. J.; Pearce, S. J.; Kiang, K. S.; Thomas, O.; de Planque, M. R. R.; Chong, H. M. H.; Morgan, H.; Ashburn, P. Ieee T Electron Dev 2014, 61, 2170-2176. (45) Prodromakis, T.; Liu, Y.; Constandinou, T.; Georgiou, P.; Toumazou, C. Ieee Electr Device L 2010, 31, 1053-1055. (46) Aronson, D.; Mittlernan, M. A.; Burger, A. J. Am J Med 2004, 116, 466-473.

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(47) Carvounis, C. P.; Nisar, S.; Guro-Razuman, S. Kidney Int 2002, 62, 2223-2229. (48) In University Science Books; University Science Books, 2014, p 372. (49) Baker, R. R.; Murray, R. K. In PDQ series; BC Decker: Hamilton, Ont., 2001, pp ix, 310 p. (50) Burton, J. J.; Garten, R. L. Advanced materials in catalysis; Academic Press: New York, 1977, p xii, 329 p. (51) Scholler, N.; Crawford, M.; Sato, A.; Drescher, C. W.; O'Briant, K. C.; Kiviat, N.; Anderson, G. L.; Urban, N. Clinical cancer research 2006, 12, 2117-2124.

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