An impedance-based integrated biosensor for suspended DNA characterisation
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1 An impedance-based integrated biosensor for suspended DNA characterisation Hanbin Ma, Richard W.R. Wallbank, Reza Chaji, Jiahao Li, Yuji Suzuki, Chris Jiggins and Arokia Nathan Supplementary Item Title or Caption Supplementary Figure 1 The overall structure and partial schematic of the CMOS biochip Supplementary Figure 2 Photographs of biochips before and after application of a wax passivation layer Supplementary Figure 3 Schematics for the impedance-based biochip system Supplementary Method Impedance extraction method Supplementary Note Effects of the electrode gap size Supplementary Note Schematic of the equivalent circuit Supplementary Note Frequency range of the measurements Supplementary Note DNA concentration detection limit and noise 1
2 Supplementary Figure 1 The overall structure and partial schematic of the CMOS biochip b. a. c. Figure S1 Figure S1a shows a block diagram of the micro-array biosensor, including the sensing area and the peripheral circuitry. To bias the pixel circuits, a two-level calibrated current mirror was used. The first level consisted of four NMOS current mirrors (CNCMs), which were calibrated in turn using a single input current from the current input to float the circuit. Each of the CNCMs was used to calibrate four PMOS current mirrors (CPCMs, supplementary Fig. S1b), which were connected to a column line through a write switch. An operational transresistance amplifier (OTRA, Supplementary Fig. S1c) was operated in current mode at the end of each column to convert the resulting current into a voltage. 2
3 Supplementary Figure 2 Photographs of biochips before and after application of a wax passivation layer Excitation and readout Circuitry Protective wax layer Sensing pads Immersed in solution Figure S2 The sensing pads were designed to be immersed in analyte solutions to perform the measurements. The excitation circuitry and bonding wires, however, need to be protected from the solution to prevent short circuits. Candle wax was used as a protective layer. We used a plastic tube to cover the sensing pads and then poured melted wax onto the chip. After the wax had solidified, the plastic tube was removed. The chip was then viewed under a microscope to ensure that the wax had covered all the circuitry. 3
4 Supplementary Figure 3 Schematics for the impedance-based biochip system Figure S3 The CMOS biochip, including the sensing and circuitry parts. An field-programmable gate array (FPGA, Altera, APEX 20K) was used to control the working pixels. A Keithley 4200 system provided the biases. A function generator (Agilent, 33250A 80 MHz) provided the stimulus signals to the biochip. An oscilloscope (Agilent, Infiniium 54833A) was used to record the input (from the function generator) and output (from the biochip) signals. 4
5 Supplementary Method Impedance extraction method Output (v out ) a. Input (v in ) Input (v in ) Output (v out ) b. Figure S4 In electrical engineering, impedance (Z) is a measure of the overall opposition of a circuit to current 1. Impendence is a more general concept than resistance because the effects of capacitance and inductance are also taken into account. Impedance is a complex ratio of the voltage to the current in an alternating current (AC) cell and can be expressed as Z = Re(Z) + j Im(Z) (1.1) or Z = Z cos(θ) + j Z sin (θ), (1.2) where Z represents the magnitude of the impedance and θ represents the phase. 5
6 For EIS, the stimulus is an AC signal, which can be expressed by v in (t) = V in max sin (ω t + φ in ). (1.3) Here, V in max is the magnitude, ω is the angular frequency, and φ in is the phase. The corresponding result is a sinusoidal current signal (i out ) with the same frequency. This current signal is converted into a voltage, v out, by the OTRA and the load resistor R L (100 kω in this work). As such, v out can be expressed as v out (t) = i out (t) R L = V out max sin (ω t + φ out ), (1.4) where the parameters used are the same as in equation 1.3 (Fig. S4.a). Therefore, the magnitude ( Z ) and the phase (θ) of the impedance can be calculated as follows: Z = V in max I out max = V in max R L V out max (1.5) and θ = φ out φ in. (1.6) The input and output signals of the sensing system, which were recorded by an oscilloscope, were read and plotted using Matlab. The parameters of the sinusoidal input and output signals can be determined automatically using the built-in curve-fit function (Fig. S4.b). Therefore, the total impedance of the measurement system can be calculated. 6
7 Supplementary Note Effects of the electrode gap size A bipolar electrode system was designed for this CMOS biochip. The effect of electrode polarisation is the primary drawback, resulting in undesirable parasitic impedance. As shown in figure S5, the planar electrodes were isolated by SiO 2, and the fringe capacitance through the solution (shown as green dashed lines) was measured. To study the effect of the gap size on electrode performance, a finite element model for two planar electrodes was simulated using the commercial software package COMSOL Multiphysics. Conductive solution Cut line for polarization calculation Gap Electrodes, Al Insulator, SiO 2 Figure S5 Cross-section of the COMSOL model Definition of the finite element model The model was developed on the basis of the top metal specifications for the 0.35 µm CMOS process. The metal pads were made of aluminium, and the patterns were separated by SiO 2. The SiO 2 barrier was 1 µm higher than the electrode s surface. Four different gap sizes, 1 µm, 10 µm, 25 µm and 50 µm, were simulated. A 1-mA current stimulus was applied to one electrode, and the other electrode was grounded. The voltage was then measured and used to calculate the solution conductivity. The frequency of the calculation was 10 khz. To study electrode polarisation, the solution conductivity was defined as S/m for a 0.08 mg/ml NaCl solution. For the sensitivity study, a set of solutions with different conductivities, 1.76 S/m, 0.9 S/m, 0.48 S/m, S/m, S/m, S/m, S/m and S/m, corresponding to different NaCl concentrations from 10 mg/ml to 0.08 mg/ml, were used. Electrode polarisation study Figure S6 depicts the simulated electrode polarisation over the electrode s surface (cut-line for polarisation calculation is labelled in figure 1 in red). The average normalised polarisation values for different electrode gap sizes were plotted. The error bars represent the standard deviation. It is evident that the polarisations of wider gaps (25 µm and 50 µm) were much smaller than those of narrower gaps (1 µm and 10 µm). Sensitivity study The sensitivity of the impedance measurements for different gap sizes was also simulated using the same model. Figure S7 illustrates the relationship between the gap size and the impedance measurement sensitivity. Sensitivity was defined as the slope of the impedance magnitude as a function of concentration. For all four sets of data, the slope was the same, These results indicate that the electrode gap size does not affect the sensitivity of the impedance measurements. In summary, a larger gap size results in smaller electrode polarisation, and the gap size does not affect the sensitivity. 7
8 Because the cost of an integrated chip is directly related to the chip size, it is not practical to design on-chip electrodes with large gaps. The 25-µm gap size used in this work is therefore a compromise between cost and performance. Figure S6. The relationship between electrode gap size and normalised polarisation. Figure S7 Calculated impedance results for different electrode gap sizes. The solution conductivities of NaCl solutions with different concentrations were measured. 8
9 Supplementary Note Schematic of the equivalent circuit The equivalent circuit is shown in Figure S8. R sol and C sol represent the solution resistance and capacitance, respectively. These parameters model the conductivity and dielectric properties of the bulk solution. At the solution-electrode interface, a double layer capacitance C dl forms owing to the applied electric field in the solution, and the polarisation behaviour can be modelled as the polarisation resistance R p. In addition, R par and Z par are included to simulate the effects of parasitic components associated with the measurement conditions. These components stem from metal interconnections, bonding wires and other parasitic elements in the chip. Figure S8 Equivalent circuit model. The test samples were saline solutions with concentrations ranging from mg/ml to 5.0 mg/ml. Figure S9 shows an example of the collected data (5 mg/ml NaCl solution) and the simulation results from 39.8 khz to 1 MHz. The impedance curve below 39.8 khz was affected by parasitic components and was thus not considered. Figure S10 presents a plot of the extracted value of the solution resistance as a function of concentration. This plot shows that the solution resistance has a strong linear correlation (R 2 = ) with the solution concentration. 9
10 Figure S9 Collected data (5 mg/ml NaCl solution) and simulation results. Figure S10 Extracted solution resistance as a function of concentration plotted on a log-log scale. 10
11 Supplementary Note Frequency range of the measurements In general, there are two factors that limit the operating frequency: the sample being tested and the CMOS transimpedance amplifier. The latter does not limit the operating frequency. The system measurement bandwidth was 1.3 MHz. The frequency limitation in our system stems from the sample being analysed. The frequency range for impedance spectroscopy can be divided into three regions based on the dominant factors 2. These factors include electrode polarisation, electrode size, surface modification and electrolyte characteristics. This information is presented in Figure S11. Charge transfer, diffusion and absorption at the electrode-solution interface are primarily observed at lower frequencies (< 100 Hz). The high frequencies (> 10 khz) are suitable for quantifying the electrical properties of the electrolyte, although there is no clear frequency demarcation. The impedance characteristics depend on many variables, such as the electrode materials, geometric area, surface modification and electrode potential. Fig. S11 Applications of and primary factors affecting bio-electrode impedance (adapted from reference 2 ). To determine the appropriate frequency range for impedance measurements of our CMOS system, we connected the system with on-chip electrodes to a commercial electrochemical impedance analyser (Solartron SI 1260). The measured equivalent impedance stems from the solution-electrode impedance plus the parasitic resistance and capacitance associated with the CMOS system. A series of saline test solutions with concentrations from mg/ml to 10 mg/ml and DI water were used. For the measurements, we applied the appropriate AC amplitude and swept the frequency from 1 khz to 1 MHz with 10 points per decade. Fig. S12 shows the Nyquist plots of impedance of the saline solutions with different concentrations and DI water. The impedance curves for different saline concentrations were easily distinguished in region II, where the frequency ranges from 39.8 khz to khz. These curves could be distinguished because the solution resistance dominates in this frequency range. An inductive loop occurs at lower frequencies (<39.8 khz) due to the parasitic circuit components. The difference in the impedance decreased above khz because the solution resistance is less dominant in this region. 11
12 Figure S13 presents the data in Figure S12 in terms of the normalised frequency-dependent sensitivity of the measurement system. The peak sensitivity was khz, and the sensitivity dropped significantly below 39.8 khz and above khz. Figure S12 Nyquist plot of the electrochemical impedance of saline solutions (concentration range: mg/ml to 10 mg/ml) and deionised water. Normalised Sensitivity I II III Frequency (khz) Figure S13 Normalised sensitivity of the measurement system extracted from the slope of the characteristic curve (impedance magnitude Z vs concentration on a log-log scale, Figure 2.1) using a linear regression algorithm. All values were normalised with respect to the peak sensitivity at khz. 12
13 Supplementary Note DNA concentration detection limit and noise To determine the detection limit for DNA concentration using on-chip impedance measurements, a DNA sample was dissolved in diluted PBS (phosphate-buffered saline, 1 unit standard PBS in 99 units of DI water) to increase the solution conductivity. DNA samples with low concentrations were prepared using the dilution procedures described in the online methods section. High frequency Low frequency Estimated noise floor Figure S14 Figure S14 presents the results of the impedance measurements for low-concentration DNA samples (1.2 ng/µl and 2.4 ng/µl) and the PBS solvent. At higher frequencies, the difference in impedance between the 1.2 ng/µl DNA sample and the solvent was less than 10 kω. Although we were still able to distinguish the two solutions from the background, the experiment yielded a rough estimate of the detection limit of DNA concentration of approximately 1 ng/µl, which is a two-fold improvement in over the limit of detection of the commercial Nanodrop spectrophotometer (2.0 ng/µl) 3. The resolution of the CMOS impedance spectroscopy chip is ultimately limited by electronic noise in the integrated readout and amplification circuitry. We measured the output noise with DI water, which was used as a solvent in this work. The noise floor of the measurement system was measured under wide bandwidth conditions. We obtained the RMS noise voltage with a zero excitation input and peak-to-peak 500 mv sinusoid voltage excitation (RMS 353 mv) and then calculated the signal-to-noise ratio, as shown in Table S5. Output noise Output signal SNR DI water 0.7 mv rms 10.6 mv rms 23.6 db Table S5. Signal-to-noise ratio at the output. 13
14 Although thermal noise is common over all stimulus frequencies, flicker noise dominates at low frequencies. As shown in Figure S14, as the DNA concentration decreased, leaving just the solvent, the impedance approached the noise floor of the system, qualitatively shown by the dashed line. Several approaches can be employed to reduce the noise floor of the impedance measurements, e.g., the transistors in the circuit can be designed to reduce flicker noise. The flicker noise (also known as 1/f noise) in CMOS transistors can be described using the following general relation 4 : < v n 2 >/f = K C ox W L (V 2 /Hz). (2.1) Here < v n 2 >/f represents the averaged noise voltage power spectral density, K and C ox are process-dependent parameters, and W L is the transistor channel area (width length). Because there is no spatial resolution requirement for the sensing pixel in biological sensing, it is possible to scale up the transistor area (W L) to reduce noise (as seen by equation 2.1). There are also other pixel design techniques that could reduce the flicker noise effectively, such as pulsed biasing 5. Alternatively, increasing the input stimulus voltage could improve the signal-to-noise ratio. 14
15 References 1. Grimnes, S. & Martinsen, Ø. G. Bioimpedance and Bioelectricity Basics. Bioimpedance Bioelectricity Basics 488 (Academic Press: 2000).at < Martinsen/dp/ > 2. Duan, Y. Factors determining and limiting the impedance behavior of implanted bioelectrodes. Smart Materials 4235, (2001). 3. ThermoSCIENTIFIC NanoDrop Products Specifications. (2012).at < 4. Ziel, A. Van Der Noise in Solid State Devices and Circuits. (Wiley: 1986). 5. Chaji, G., Nathan, A. & Pankhurst, Q. Merged phototransistor pixel with enhanced near infrared response and flicker noise reduction for biomolecular imaging. Applied Physics Letters 3 5 (2008).doi: /
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