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Development of Microelectrode Arrays Using Electroless Plating for CMOS-Based Direct Counting of Bacterial and HeLa Cells Kiichi Niitsu, Member, IEEE, Shoko Ota, Kohei Gamo, Student Member, IEEE, Hiroki Kondo, Masaru Hori, and Kazuo Nakazato
Abstract—The development of two new types of high-density, electroless plated microelectrode arrays for CMOS-based high-sensitivity direct bacteria and HeLa cell counting are presented. For emerging high-sensitivity direct pathogen counting, two technical challenges must be addressed. One is the formation of a bacteria-sized microelectrode, and the other is the development of a high-sensitivity and high-speed amperometry circuit. The requirement for microelectrode formation is that the gold microelectrodes are required to be as small as the target cell. By improving a self-aligned electroless plating technique, the dimensions of the microelectrodes on a CMOS sensor chip in this work were successfully reduced to 1.2 2.05 . This is 1/20th of the smallest size reported in the literature. Since a bacteria-sized microelectrode has a severe limitation on the current flow, the amperometry circuit has to have a high sensitivity and high speed with low noise. In this work, a current buffer was inserted to mitigate the potential fluctuation. Three test chips were fabricated using a 0.6CMOS process: two with 1.2 2.05 (1024 1024 and 4 4) sensor arrays and one with 6square (16 16) sensor arrays; and the microelectrodes were formed on them using electroless plating. The uniformity among the 1024 1024 electrodes arranged with a pitch of 3.6 4.45 was optically verified. For improving sensitivity, the trenches on each microelectrode were developed and verified optically and electrochemically for the first time. Higher sensitivity can be achieved by introducing a trench structure than by using a conventional microelectrode formed by contact photolithography. Cyclic voltammetry (CV) measurements obtained using the 1.2 2.05 4 4 and 6square 16 16 sensor array with electroless-plated microelectrodes successfully demonstrated direct counting of the bacteria-sized microbeads and HeLa cells. Index Terms—Bacteria counting, CMOS, electroless plating, HeLa cells, microelectrode array, point-of-care testing.
I. INTRODUCTION OUNTING pathogens (cells, viruses, and bacteria) directly in real time with high sensitivity is useful in maintaining human health and preventing pandemics or bioterrorism
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Manuscript received January 16, 2015; revised June 15, 2015; accepted September 06, 2015. Date of publication November 06, 2015; date of current version November 24, 2015. This work was supported by JSPS KAKENHI Grants 20226009, 25220906 and 24760266, SCOPE Grants 121806006 and 152106004, JST/PRESTO, and VDEC. This paper was recommended by Associate Editor S. Sonkusale. K. Niitsu is with Nagoya University, Nagoya 464-8601, Japan, and also Japan Science and Technology Agency (JST) PRESTO Researcher, Saitama 332-0012, Japan (e-mail:
[email protected]). S. Ota was with Nagoya University, Nagoya 464-8601, Japan. K. Gamo, H. Kondo, M. Hori, and K. Nakazato are with Nagoya University, Nagoya 464-8601, Japan. Color versions of one or more of the figures in this paper are available online at http://ieeexplore.ieee.org. Digital Object Identifier 10.1109/TBCAS.2015.2479656
[1] (Fig. 1). However, conventional counting methods with high sensitivity such as the polymerase chain reaction [2] and culture-based biochemical testing [3] are time consuming (usually a few days at minimum). For example, the PCR usually takes a few hours and the culture-based testing takes usually a few days at minimum. These methods also require controlled environments and well-trained staff. On the other hand, conventional rapid approaches such as immuno-chromatography [4] and ATP bioluminescence schemes [5] are insufficient with respect to sensitivity. Flow cytometry [6] has high throughput and high sensitivity. Even though mobile flow cytometry [7] and magnetic flow cytometry [8] are remarkably developed, it requires relatively large-volume apparatus because it is based on optical measurement when comparing to electrical one. The Coulter counter [9] can also be used to achieve high throughput, however, the required apparatus is very big and expensive. As an alternative, electro-chemical sensing approaches using a microelectrode [10]–[13] have been intensely investigated as shown in Fig. 1(a). By using antibody immobilization on a CMOS-based microelectrode array [14], [15], simultaneous counting of various pathogen types can be implemented. To maximize sensitivity, the electrode size should be same as that of the target [16]. However, the sizes of conventional microelectrodes are not small enough for direct bacteria counting. For example, the size of the state-of-the-art microelectrode is 7square [12], which is much larger than the size of bacteria (2–8 ). This is insufficient for high-sensitivity direct counting as shown in Fig. 1(b). Au electrodes are commonly used in electrolyte systems because Au has the lowest ionization tendency. However, the formation of Au microelectrodes is not allowed in standard CMOS processes because Au may cause serious contamination. Thus, Au microelectrodes must be formed with a post-CMOS process. However, the conventional post-CMOS process is based on either contact photolithography or electron-beam lithography. Contact photolithography is insufficient in terms of miniaturization limit because its resolution is in several micrometer lines/spaces. In addition, it requires large and expensive production equipment. From the viewpoint of miniaturization limit, electron-beam lithography is attractive. However, electron-beam lithography also requires large and expensive production equipment. Moreover, it requires a long time and is unsuitable for mass production. Thus, cheap and self-aligned processes are needed for the development of bacteria-sized microelectrodes, as shown in Fig. 2.
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Fig. 1. Conceptual diagram of the proposed work. (a) Direct counting. (b) Size of microelectrode.
1024 electrodes arranged with a pitch of 3.6 4.45 was verified using a laser microscope. In addition, the trenches on each microelectrodes for improving sensitivity were developed and optically verified for the first time. The trench structure aids in single target capture and increases the sensitivity of the electrochemical measurement [6]. CV for the 16 16 sensor array chip with 6square electroless-plated microelectrodes successfully demonstrated the effectiveness of the platform in direct counting of HeLa cells. The size of the HeLa cells is larger than that of the bacteria. This study presents the first demonstration of pathogen detection using an electroless-plated gold microelectrode; the previous study [12] on this subject did not detect any pathogen or microbeads. Thus, the present work is useful toward developing a method for bacteria detection. In addition to the conference proceedings [30], a detailed description of the implementation of amperometry, CV measurements and direct counting of the bacteria-sized microbeads using a 1.2 2.05 4 4 microelectrode array, 2D mapping results, and 2 2 array measurement results are presented in this paper. This paper is organized as follows. The proposed formation of microelectrodes by the electroless plating is introduced in Section II. Design of two prototype chips and the measurement setup are shown in Section III. In Section IV, we present measurement results. Section V shows 2D counting results. Sections VI and VII show direct counting of the bacteria-sized microbeads and HeLa cells. Section VIII concludes this paper. II. ELECTROLESS GOLD PLATING FOR SMALL-FEATURE MICROELECTRODE FORMATION A. Formation Process of Electroless Gold Plating
Fig. 2. Comparison of the fabrication methodologies for forming gold electrode. ( means good, means medium, and means bad.)
Electroless plating is a promising technique for forming bacteria-sized Au microelectrodes because it is a cheap self-aligned process and forms trench structures. In this work, by introducing and improving an electroless plating technique, the size of microelectrodes on a CMOS sensor chip was successfully reduced to 1.2 2.05 , which corresponds to 1/20th (from 49 to 2.46 ) of the smallest size reported in literature as shown in Fig. 1(b). To verify the effectiveness of the proposed approach, three test chips were fabricated using a 0.6CMOS process: two with 1.2 2.05 sensor arrays (1024 1024 and 4 4), and one 6square sensor array (16 16). The Au microelectrodes were formed on them. The uniformity among 1024
Fig. 3 shows the flowchart and microphotograph of the electroless gold plating technique for realizing the 1.2 2.05 microelectrode array. Though the shape should be optimized for the target, the shape is not a square but a rectangle because of the design rule of CMOS fabrication in this work. The aluminum alloy on the surface of the CMOS chip is susceptible to oxidization and thus is difficult to perform electroless-plating. To achieve a small form factor, the following key points are considered. Contaminations adversely affect the adhesion and flatness of microelectrodes. Therefore, pre-cleaning is important to achieve small size and uniformity. We performed pre-cleaning using acetone, isopropyl alcohol (IPA), and de-ionized (DI) water. We have based some steps of the proposed procedure on [12], combining them with pre-processing [17] and de-smutting [18]. During the electroless plating, the chip did not require any protection. The challenge is the temperature management. Temperature management is important especially in small-feature (micrometer-order) electrode formation. Electroless-plating has an advantage in terms of duration. The thickness of the layer formed by electroless-plating is approximately 0.2 . The microelectrode was subjected to repeated measurement. CV measurement results did not change after seven times of repeated measurement. In addition, it could
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Fig. 4. Microphotograph and 3D image of the formed electroless-plated gold 2.05 ) in the 1024 1024 chip. microelectrode (1.2
Fig. 3. Flowchart and microphotograph of electroless gold plating for realizing 2.05 microelectrode. Right part is the microphotograph of the a 1.2 square electrode. formed 100-
maintain its performance after 2-minutes ultrasonic cleaning (100 W, 35 kHz). B. Trench Structure by Electroless Gold Plating A trench structure is introduced to improve sensitivity. The current which generated from the peripheral part of the microelectrode is unaffected by the target and degrades sensitivity. To minimize the current flow from the peripheral part of the microelectrode while maintaining the diameter, we introduced the trench structure. One of the previous studies [8] presented quantitative discussion of sensitivity improvement using the trench structure. The said paper reported that a trench structure (1.5- -diameter microelectrode surrounded by 1- -high wall to detect a 1- -diameter target) improves the sensitivity by a factor of four (from 8% to 30%). In addition, the trench structure mitigates the expansion of diffusion layer and allows the decrease of pitch. The trench structure was formed by an opening at the top passivation layer using standard CMOS process and electroless plating. A trench structure with gold microelectrode can be realized using electroless plating with a thickness that is thinner than the height of the top passivation layer. In the conventional contact-photolithography-based technology, realizing a bacteria-sized trench structure is difficult because making the post-CMOS passivation layer thick is difficult and a high-aspect trench formation is not practical with respect to the resolution limitation of contact photolithography. C. Verification of Electroless Gold Plating by Optical View Three test chips were fabricated in a 0.6CMOS process, with sensor arrays of 1.2 2.05 (1024 1024 and 4 4), and 6square (16 16). The Au microelectrodes
Fig. 5. Step profile of the formed electroless-plated gold microelectrode 2.05 ) in the 1024 1024 chip. (a) Microphotograph taken by (1.2 laser microscope. (b) Vertical structure taken by SEM. (c) Vertical measurement of trench structure by laser microscope.
were formed on them. Fig. 4 shows a microphotograph and 3D image of the electroless-plated gold microelectrode (1.2 2.05 ) on the 1024 1024 chip. Fig. 5 shows the cross-sectional step profile of the microelectrode. The successful formation of microelectrodes with good uniformity and flatness was verified by viewing through a hybrid laser microscope (OPTELICS HYBRID) manufactured by Lasertec Co., (Kanagawa, Japan). The trench structure for increasing the sensitivity of individual single bacteria detection was also verified as shown in Fig. 5(c). Fig. 6 shows the microphotograph and 3D image of the fabricated electroless-plated Au microelectrode (6square) in the 16 16 chip. The formation of the microelectrode (6square) was also verified. From the two types of the formed arrays, the scalability of the microelectrode’s size from 1.2 2.05 to 6square using the proposed procedure was confirmed. The following chemicals and reagents were used in the electroless plating process: RONAMAX SMT-115 purchased from Dow Chemical Japan Co., (Tokyo, Japan); OROMERSE SO, TECHNI SBZ CONDITIONER purchased from Technic Japan Co., Ltd. (Osaka, Japan); nitric acid purchased from Yoneyama
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Fig. 6. Microphotograph and 3D image of the formed electroless-plated gold square) in the 16 16 chip. microelectrode (6-
Yakuhin Kogyo Co., Ltd. (Osaka, Japan); and sodium hydroxide purchased from Sigma-Aldrich (St. Louis, Missouri, USA). D. Scaling Limit of Electroless Gold Plating A smaller size of microelectrodes enables detection of smaller pathogens such as viruses (20–970 nm). In addition, imaging-based counting (not direct counting) that does not suffered from the limited fill factor can be feasible. However, two types of scaling limitations for electroless plating exist: electrode-size and electrode-gap limitations. The scaling of electrode size depends on reaching an electroless-plating solution and on the CMOS fabrication. Reaching an electroless-plating solution is essential for electroless plating. If the electroless-plating solution cannot reach the original electrode before the electroless plating, the reaction involved in electroless plating cannot start. CMOS fabrication can be scaled using the current scaled CMOS technology. In the 0.6CMOS technology used in this work, 1.2 2.05 is minimum feature size because of the design constraints (design rule check of the CMOS fabrication process). Using scaled CMOS technology such as the 14-nm CMOS [19] allows the development of sub-micrometer microelectrodes. Since RF-CMOS integration requires highly-integrated top metal layer for small-formfactor high- on-chip inductor, top metal layer is also being scaled down. The literature reported that minimum pitch of top metal layer was scaled down to 0.6 in 32 nm CMOS [20]. The electrode-gap scaling is limited by the self-terminating property of the electroless gold plating. Previous work in the nanotechnology field [21] has shown that a 5-nm gap can be developed by electroless gold plating. In addition, the gap can be adjusted by introducing a molecular ruler [22]. E. Fill Factor of Electroless-Plated Microelectrode Array In direct bacterial counting using a 2D microelectrode array, a limited fill factor decreases the counting accuracy. Limited fill factor is caused by the gap between microelectrodes. The gap of a conventional contact-photolithography-based microelectrode
Fig. 7. Circuit structure of the microelectrode array.
array is limited by the resolution of the contact photolithography, and its minimum value is greater than several micrometers. The gap of the proposed electroless-plated microelectrode array is limited by the self-termination property, and its minimum value is approximately 5 nm [21]. Furthermore, the trench structure is beneficial to improve the counting accuracy as well as the sensitivity. The trench structure usually makes the gap assume a convex shape as shown in Fig. 5. By introducing this shape, the pathogen cannot stay at the gap and will fall inside the trench due to gravity. Therefore, the electroless-plating technique has an advantage from the viewpoint of the fill factor. To improve the counting accuracy in electroless-plated microelectrode arrays with a trench structure, introduction of magnetic cell manipulation using an on-chip coil [23] can be one of the effective solutions. By immobilizing the magnetic beads on pathogens and generating magnetic field using an on-chip coil, pathogens can be accumulated inside the trench. As a result, the degraded counting accuracy due to the limited fill factor can be mitigated. III. TEST-CHIP DESIGN AND MEASUREMENT SETUP FOR AMPEROMETRY A. Test Chip Design Since the current flow on bacteria-sized microelectrode is severely limited, the amperometry circuit needs to be low noise while maintaining high speed. In this work, some circuit techniques were implemented to address this issue. The circuit architecture is shown in Fig. 7. The 1.2 2.05 1024 1024 microelectrode array chip does not employ the current buffer circuits. The 1.2 2.05 4 4 microelectrode array chip employs a current integrator [24], [25] instead of the - converter. Amperometry using a microelectrode requires several tens of seconds for stabilization of the signal (current). Therefore, the measurement requires an unacceptably long time if we measure the sensors one by one. Thus, in order to obtain an accurate data, amperometry using a microelectrode requires a fixed electrode potential and a steady-state current must flow from the electrode continuously during measurement.
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Fig. 8. Timing diagram of the signals of the selection switch.
As a countermeasure to this issue, we employ selection switch [26] for the emerging fixed electrode potential, as in shown Fig. 7. For a fixed electrode potential, there are two switches (switch 1 and switch 2) per electrode. Without connecting the – converter, the electrodes are fixed at constant electrode potential. Since a constant steady-state current flows continuously through electrodes with the fixed potentials, it is possible to reduce the time to reach steady state and thus the total measurement time. The switch operation is as follows. When electrode is measured, switches 1A and 2B remains closed, and switches 2A and 1B are open. Switches 2A and 1B are close, and switches 1A and 2B are opened when we measure electrode B. As such, it is possible to perform a high-speed measurement when a constant current flows by maintaining all electrode potentials at constant values. Fig. 8 shows the selection switch signals. By turning off switch 2 after turning on switch 1, the potential of the electrode is always fixed. Fig. 9 shows a schematic of the proposed sensor cell circuit for amperometry. The sensor cell circuits for amperometry [27] were implemented in a 16 16 array chip that consists of 16 16 blocks. One block consists of four different sensors: amperometric, potentiometric, impedimetric, and simple electrode sensors. In this work, an amperometric sensor was used for performing CV measurements. The selection switch may cause potential variations in the electrode during the switching operation since it is directly connected to the electrode. The potential variation may cause unstable amperometric measurement, which may lead to damage of the microelectrodes. As a countermeasure to this, we implemented a current buffer circuit between the electrode and the selection switch for reducing the voltage variation, as depicted in Fig. 7. By duplicating the current flowing through the electrode, the effect of switching on the measurement system is reduced. As mentioned in Fig. 9, the current buffer circuit for amperometry consists of a current mirror to copy the current, source followers, a differential amplifier, MOS diodes, and a level shift source follower. The groups of P1–P4, P5–P6, N1–N4, {N5, N11, N12}, and N8–N10 use the same size transistors. The current mirrors of {P1, P2} and {N4, N3} are set to copy the drain current from P2 to P1 and N4 to N3. The source follower of {N2, P4} is set to maintain the electrode voltage at a constant value. The differential amplifier of {P5, P6, N5, N6, N9} creates the gate potential of the source follower ( ). N1 and P3, which are MOS diodes, provide larger output imped-
Fig. 9. Schematic of the sensor cell circuit (current buffer and selection switch) for amperometry.
ances. The switching circuit of {P7, N7} selects one electrode from the array of electrodes. The source follower of {N12, N8, N11, N10} increases operating range. Fig. 10 shows the schematic of the amperometric sensor array. Sides of the array have decoders ( decoder and decoder) and address buffers, respectively. The address buffers are used to hold the address signal and are reset by the CAS for the address and the RAS for the address. The and decoders convert the signals from the address buffer and set the voltage of the word line, which we want to read as low. The sensor array is measured as follows: the addresses are reset by the RAS and CAS. The address of the specific electrode sets the input address. The address buffer of the axis maintains the input address. The address of the specific electrode sets the input address. The address buffer of the axis maintains the input address. The switch between the current circuit and the – converter is set to on. The off-chip analogto-digital converter (ADC) measures the potential of . When the next electrode is measured, measurement starts from the step that the address buffer of axis maintain the input address. When the line of the electrode ( address) is changed, measurement starts from the step of the reset. Fig. 11 shows a schematic of the – converter. Fig. 11(a) shows that the current ( ) is amplified and is converted into a voltage ( ) by the resistor ( ) in the – converter. The potential of is ( , 10, 100, 1000). The amplified current is generated by the current mirror in Fig. 11(b). B. Measured Results of the Test Chip The measured results of the test chip are summarized in this subsection. The detail can be found in previous literature [13].
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Fig. 12. Measured input-output characteristic of the current buffer [23].
Fig. 10. (a) Microphotograph of the 16 of the sensor array for amperometry.
16 sensor array. (b) Block diagram
Fig. 13. Measured input-output characteristic of the 100, 1000).
Fig. 11. Schematic of the - converter. (a) Block diagram of the converter. (b) Schematic diagram of the current mirror in the - converter.
Fig. 12 shows the measured input-output characteristic of the current buffer shown in Fig. 9. The input signal was supplied by HP4142B (Keysight Technologies). The measured range of
converter (
, 10,
the input current is from 3 pA to 2 , which agrees well with the SPICE simulation results. The linear region is wide enough for our application. The measured current gain is one, and the measured offset current is approximately 2 nA. The leakage of the switches between the current buffer and the - converter is sufficiently small; thus, they can be ignored. The simulated equivalent input-referred current noise integrated from 10 Hz to 10 kHz is 21.8 pA. Fig. 13 shows the measured input-output characteristic of the - converter shown in Fig. 9. Good linearity is verified. The log-log plot of the measured data at is shown in Fig. 14. Because of process variation, the slope changes at the boundary (the input current is zero).
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Fig. 14. Measured log-log plot of the input-output characteristic of the converter ( ).
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Fig. 16. Measured CV of 1.2
2.05
4
4 microelectrode array.
IV. MEASUREMENT RESULTS A. CV of a 1.2
2.05
4
4 Microelectrode Array
To confirm the electrochemical property of the electrolessplated gold microelectrode, we performed CV measurement for a 1.2 2.05 4 4 microelectrode array using 20-mM and . Since this measurement is feasibility study, an external potentiostat was used for conservative approach. Fig. 16 shows the measured result. The measured peak current was 3.75 nA. The theoretical formula for peak current is expressed as (1)
Fig. 15. Measurement setup for CV measurements and HeLa cell counting.
C. Measurement Setup Fig. 15 shows the measurement setup. A semiconductor parameter analyzer (HP 4142B) and a potentiostat were utilized for the CV measurement. Micro controller unit (MCU, STM32F103VET6, ST Microelectronics) was implemented for signal processing. The clock frequency of the MCU is 72 MHz. A 12-bit 1- ADC is implemented in the MCU. This measurement system can achieve rapid bacterial counting (approximately 1 s) even with a 1024 1024 array. The following chemicals and reagents were used in amperometry: phosphate-buffered saline (PBS, pH 7.4) purchased from Wako Pure Chemical Industries, Ltd. (Osaka, Japan), sodium sulfate, and potassium hexacyanoferrate (II) trihydrate purchased from Kanto Chemical Co., Inc. (Tokyo, Japan).
where is the number of electrons, is the Faraday constant, is the diffusion constant, is the density of the reactant, and is the radius of the microelectrode [11]. From (1), the theoretical value of the peak current in the 1.2 2.05 microelectrode is 4.05 nA. Thus, the measured values are consistent and slightly less than the theoretical value. Two possible reasons could have caused this decrease. One is that the solution cannot reach the edge of the trench of the microelectrodes and the effective area decreases. Another is that the trench structure for sensitivity improvement reduces the absolute value of the peak currents, as reported in [10]. B. CV of 6-
Square 16
16 Microelectrode Array
Fig. 17 shows the measured CV using 10 mM and . Fig. 17(a) shows the results of the median of the 16 16 array when the current was measured seven times.
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A previous work [23] reported that the standard deviation value for the conventional contact-photolithography-based CMOS amperometric sensor array with a 10 -square microelectrode is 3.94 nA when the average peak current is 15.4 nA. Current buffers with identical design were implemented in the previous and present works. Thus, the proposed electroless-plated microelectrodes have an advantage from the viewpoint of performance variation. V. 2D COUNTING USING ARRAY SENSORS We have confirmed the principle of operation of 2D counting by using the plated electrode array. Fig. 18(a) shows a conceptual diagram of the principle of 2D counting. Silicone coated part of the electrode array, as shown as Fig. 18(b), and the electrode array was measured by CV to confirm the principle. Fig. 18(d) shows a clear difference in the electrodes with and without silicone, as the electrode with silicone does not allow a current to flow. Comparing Fig. 18(b) with Fig. 18(c), we find that silicone form 2D counting is possible by the chip. VI. DIRECT COUNTING OF BACTERIA-SIZED MICROBEADS
Fig. 17. Measured results of cyclic voltammetry (CV). (a) Results of seven iterations. (b) Histogram of the measured peak current.
The measured current is almost identical during all seven measurements. From the measurements, we found that the electroless-plated microelectrode was robust enough to be used several times. The peak currents from the first, third, and seventh measurements were . The drift is less than 0.1 nA. From (1), the theoretical value of the peak current in the 6square microelectrodes is 7.55 nA. The reason of the decrease from the theoretical value is consider as well as the case of 1.2 2.05 4 4 microelectrode array in the previous subsection. Fig. 17(b) shows the histogram of the measured peak currents in the first measurement. All 16 16 electrodes functioned well. The histogram in all 16 16 electrodes of the limiting current forms a normal distribution and its standard deviation is 0.305 nA, where the average peak current is 5.33 nA. The distribution comes from the variation in both the area of the microelectrodes and the current buffers. Fig. 5 shows that the surface roughness of the electroless-plated microelectrodes is approximately 0.1 . Thus, we assumed that the variation in the area of the microelectrodes is small and the distribution might have come mainly from the current buffers.
To evaluate the feasibility of direct bacteria counting using the electroless-plated microelectrode, the direct counting of bacteria-sized microbeads was performed using 1.2 2.05 4 4 microelectrode array with 10-mM and . To imitate the bacteria, the magnetic microbeads (Dynabeads MyOne Streptavidin C1, Life Technologies Corp., ) were used. By attaching a neodymium diameter is 1 magnet to the bottom of the ceramic package, magnetic microbeads are immobilized at the surface of the chip. Fig. 19 shows the measured results of the direct counting. The CV’s peak current was reduced from 2.54 nA to 1.83 nA. The ) is sufficient to maintain signal-todifference of 0.71 nA ( noise ratio (SNR) under drift (less than 0.1 nA) and the input-referred noise (21.8 pA). The direct counting of microbeads was successfully verified. This results show the feasibility of the direct bacteria counting by the proposed method. As well as qualitatively, the result is quantitatively reasonable. The study [8] reported that the sensitivity was improved to by the trench structure where 1.5- -difrom ameter microelectrode surrounded by 1- -high wall to detect a 1- -diameter target. This insight agrees well with the measured results in this study. Therefore, we assumed that the meais owing to the trench structure. sured sensitivity of VII. DIRECT COUNTING OF HELA CELLS A. Introduction and Concept To verify the effectiveness of the proposed platform, direct counting of HeLa cells was demonstrated by using the 16 16 sensor chip. The HeLa cell line [28] is an immortal cell line, which is widely and commonly used in scientific research [29]. The line is derived from human cervical cancer. HeLa cells were chosen owing to their remarkable durability. Fig. 20 illustrates the conceptual diagram of the principle of direct counting of HeLa cells. The presence of HeLa cells on the microelectrode reduces the current from the counter electrode (CE) to the
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Fig. 20. Conceptual diagram of the principle of direct counting of HeLa cells.
Fig. 18. 2D counting. (a) Conceptual diagram of 2D counting.
To verify the bare electrode behavior, CV with a 10 mM without was performed. Because HeLa cells shrink after death, they must be kept active during measurement. To minimize influence of electrolytes on the activity of HeLa cells, was left off. After verifying the microelectrode with CV, 40 of HeLa cells (8.1 ) solution in PBS was dripped and incubated for 2 h at room temperature. This is done to allow the cells to settle on the microelectrodes. Measurements were taken after drawing out the unnecessary solution (PBS with redundant cells that did not settle on the microelectrodes). Fig. 20 shows how to determine whether the microelectrode is with or without a HeLa cell. Because the size of one HeLa cell is about 10 , it can cover the entire microelectrode. When a cell is attached to the microelectrode, the current is blocked by the cell, reducing the amount of current. By detecting this change, we tried to determine whether a cell is on the microelectrode or not. B. Results of Direct Counting of HeLa Cells
Fig. 19. Measured CV of 1.2 2.05 without and with the bacteria-sized microbeads.
4
4 microelectrode array
working electrode (WE). Thus, by monitoring the difference in the currents, HeLa cells can be counted directly.
Fig. 21 demonstrates the microphotograph of the CMOS sensor surface with well-distributed HeLa cells. Though HeLa cells gather at the center of the chip, the measurement is unaffected. Figs. 22 and 23 show the measured current as a function of RE’s potential without and with HeLa cell. The difference between (a) and (b) is their positions of the microelectrodes. The curves shown in the middle of Figs. 22 and 23 are raw data obtained from the chip. They contain the DC offset current due to the output buffer circuit. There is a difference among the peak currents of CV using in Figs. 22 and 23 and that in Fig. 17(a). This difference is due to the different supporting electrolyte [PBS in Figs. 22 and 23, in Fig. 13(a)]. In Fig. 22, the peak current without HeLa cells decreases by 36% and 35%, since the PBS employed for cells settle on the microelectrode and reduce the concentration of . On the other hand, in Fig. 23, the peak currents with the presence of HeLa cells decrease by 52% and 50%. The difference in decrement (36% and 35% versus 52% and 50%) is due to the presence of HeLa cells. Thus, the measured results indicate that the proposed sensor platform can detect HeLa cells directly with high sensitivity. As the lower limit of detection of the implemented circuit is approximately 3 pA [26]. Even the microelectrode with
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Fig. 21. Microphotograph of the CMOS sensor surface with distributed HeLa cells.
Fig. 23. Measured current as a function of RE’s potential with HeLa cells.
decreases by 44%, and the peak current of the right electrodes with a cell decreases by 60%. By determining the threshold between the left electrodes (without a cell) and the right electrodes (with a cell) and identifying the peak current, we can determine if the electrode is a “with” or “without” cell using this chip. The obtained cell count from the unbiased direct electrical detection agree with that from the microscope image. To verify the effectiveness of the current decrease in Fig. 23, we performed CV using 10-mM and PBS. First, . Next, we performed CV using 2 ml of 10-mM we dropped 40 of PBS onto the chip and performed CV again using 2 ml of . Fig. 25(b) shows the results of the median current of a 12 12 array in the red frame in Fig. 25(a). The results of the median of CV do not include the upper-left sensors because they did not function owing to an air bubble. The peak CV current decreases because PBS reduced the concentration of . C. Summary of System Specification Fig. 22. Measured current as a function of the RE’s potential without HeLa cells.
the smallest peak current in Fig. 17(b) can detect the peak-current difference of greater than 200 pA owing to the presence of HeLa cells. Fig. 24 shows the results for the 2 2 array. The results are obtained without any post processing such as a clustering algorithm. The peak current of the left electrodes without a cell
The total system specification can be summarized from the measurement results. In direct pathogen counting, differentiating the pathogens from the drift and debris is essential. From the CV measurement results, we found that the drift was approximately less than 0.1 nA. The signal from the HeLa cell counting was greater than 0.45 nA (15% of 3 nA). Thus, sufficient SNR can be obtained. Further, because the electroless-plated microelectrodes are durable and washable, the debris can be removed by washing.
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array using electroless plating for CMOS-based high-sensitivity direct bacteria counting. By adopting and improving the electroless plating technique, the size of the microelectrodes on the CMOS sensor chip can be reduced to 1.2 2.05 . The uniformity among 1024 1024 array and trench structure for improving sensitivity were verified. CV measurements with 1.2 2.05 4 4 sensor array demonstrated the direct counting of the bacteria-sized microbeads while achieving the enhanced sensitivity using the trench structure. The measurements taken with the 16 16 sensor array chip with 6square electroless-plated microelectrodes successfully demonstrated well-uniformed ( ) CV and direct counting of HeLa cells. REFERENCES
Fig. 24. The result of 2D direct HeLa cell detection with microelectrode array (2 2).
Fig. 25. Measurements for CV. (a) 2D map of current peak. (b) Results of the median value of the 12 12 array.
VIII. CONCLUSION This paper has reported the development of the bacteria-sized (1.2 2.05 ) 1024 1024 microelectrode array and HeLa cell counting with 6 square 16 16 microelectrode
[1] D. Ivnitski, I. Abdel-Hamid, P. Atanasov, and E. Wilkins, “Biosensors for detection of pathogenic bacteria,” Biosens. Bioelectron., vol. 14, no. 7, pp. 599–624, Oct. 1999. [2] K. E. Templeton, S. A. Scheltinga, M. F. C. Beersma, A. C. M. Kroes, and E. C. J. Claas, “Rapid and sensitive method using multiplex realtime PCR for diagnosis of infections by influenza A and influenza B viruses, respiratory Syncytial virus, and Parainfluenza viruses 1, 2, 3, and 4,” J. Clin. Microbiol., vol. 42, no. 4, pp. 1564–1569, Apr. 2004. [3] T. K. W. Ling, Z. K. Liu, and A. F. B. Cheng, “Evaluation of the VITEK 2 system for rapid direct identification and susceptibility testing of Gram-negative Bacilli from positive blood cultures,” J. Clin. Microbiol., vol. 41, no. 10, pp. 4705–4707, Oct. 2003. [4] H. Arai, B. Petchclai, K. Khupulsup, T. Kurimura, and K. Takeda, “Evaluation of a rapid immunochromatographic test for detection of antibodies to human immunodeficiency virus,” J. Clin. Microbiol., vol. 37, no. 2, pp. 367–370, Feb. 1999. [5] E. W. Chappelle and G. V. Levin, “Use of the firefly bioluminescent reaction for rapid detection and counting of bacteria,” Biochem. Med., vol. 2, no. 1, pp. 41–52, Jun. 1968. [6] O. Fornas, J. Garcia1, and J. Petriz, “Flow cytometry counting of CD34+ cells in whole blood,” Nature Med., vol. 6, no. 7, pp. 833–836, Jul. 2000. [7] Q. Wei, H. Qi, W. Luo, D. Tseng, S. J. Ki, Z. Wan, Z. Göröcs, L. A. Bentolila, T.-T. Wu, R. Sun, and A. Ozcan, “Fluorescent imaging of single nanoparticles and viruses on a smart phone,” ACS Nano, vol. 7, no. 10, pp. 9147–9155, Sep. 2013. [8] M. Helou, M. Reisbeck, S. F. Tedde, L. Richter, L. Bär, J. J. Bosch, R. H. Stauber, E. Quandt, and O. Hayden, “Time-of-flight magnetic flow cytometry in whole blood with integrated sample preparation,” Lab. Chip., vol. 13, no. 6, pp. 1035–1038, Mar. 2013. [9] J. Zhe, A. Jagtiani, P. Dutta, J. Hu, and J. Carletta, “A micromachined high throughput Coulter counter for bioparticle detection and counting,” J. Micromech. Microeng., vol. 17, no. 2, pp. 304–313, Jan. 2007. [10] Y. Ishige, Y. Goto, I. Yanagi, T. Ishida, N. Itabashi, and M. Kamahori, “Feasibility study on direct counting of viruses and bacteria by using microelectrode array,” Electroanal., vol. 24, no. 1, pp. 131–139, Jan. 2012. [11] Y. Chen, C. C. Wonga, T. S. Puia, R. Nadipallib, R. Weerasekeraa, J. Chandrana, H. Yub, and A. R. A. Rahmana, “CMOS high density electrical impedance biosensor array for tumor cell detection,” Sens. Actuators B, Chem., vol. 173, pp. 903–907, Oct. 2012. [12] S. Hwang, C. N. LaFratta, V. Agarwal, X. J. Yu, D. R. Walt, and S. Sonkusale, “CMOS microelectrode array for electrochemical lab-ona-chip applications,” IEEE Sensors J., vol. 9, no. 6, pp. 609–615, Jun. 2009. [13] L. Li, X. Liu, W. A. Qureshi, and A. J. Mason, “CMOS amperometric instrumentation and packaging for biosensor array applications,” IEEE Trans. Biomed. Circuits Syst., vol. 5, no. 5, pp. 439–448, Oct. 2011. [14] Q. Xuea, C. Biana, J. Tonga, J. Suna, H. Zhanga, and S. Xia, “CMOS and MEMS based micro hemoglobin-A1c biosensors fabricated by various antibody immobilization methods,” Sens. Actuators A, Phys., vol. 169, no. 2, pp. 282–287, Oct. 2011.
618
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[15] J. Cooper, N. Yazvenko, K. Peyvan, K. Maurer, C. R. Taitt, W. Lyon, and D. L. Danley, “Targeted deposition of antibodies on a multiplex CMOS microarray and optimization of a sensitive immunoassay using electrochemical detection,” PLoS ONE, vol. 5, no. 3, p. e9781, Mar. 2010. [16] S. K. Arya, K. C. Lee, D. B. Dah’alan, Daniel, and A. R. A. Rahman, “Breast tumor cell detection at single cell resolution using an electrochemical impedance technique,” Lab on a Chip, vol. 12, no. 13, pp. 2362–2368, Jul. 2012. [17] M. Datta, S. A. Merritt, and M. Dagenais, “Electroless remetallization of aluminum bond pads on CMOS driver chip for flip-chip attachment to vertical cavity surface emitting lasers (VCSEL’s),” IEEE Trans. Compon. Packag. Technol., vol. 22, no. 2, pp. 299–306, Jun. 1999. [18] Kanigen Tech. Rep. No. 9 (in Japanese) [Online]. Available: http:// www.kanigen.co.jp/file/report9.pdf [19] S. Natarajan, M. Agostinelli, S. Akbar, M. Bost, A. Bowonder, V. Chikarmane, S. Chouksey, A. Dasgupta, K. Fischer, Q. Fu, T. Ghani, M. Giles, S. Govindaraju, R. Grover, W. Han, D. Hanken, E. Haralson, M. Haran, M. Heckscher, R. Heussner, P. Jain, R. James, R. Jhaveri, I. Jin, H. Kam, E. Karl, C. Kenyon, M. Liu, Y. Luo, R. Mehandru, S. Morarka, L. Neiberg, P. Packan, A. Paliwal, C. Parker, P. Patel, R. Patel, C. Pelto, L. Pipes, P. Plekhanov, M. Prince, S. Rajamani, J. Sandford, B. Sell, S. Sivakumar, P. Smith, B. Song, K. Tone, T. Troeger, J. Wiedemer, M. Yang, and K. Zhang, “A 14 nm logic technology featuring 2nd-generation FinFET, air-gapped interconnects, self-aligned double patterning and a 0.0588 SRAM cell size,” in Proc. IEEE Int. Electron Device Meeting, Dec. 2014, pp. 3.7.1–3.7.3. [20] C.-H. Jan, M. Agostinelli, H. Deshpande, M. A. El-Tanani, W. Hafez, U. Jalan, L. Janbay, M. Kang, H. Lakdawala, J. Lin, Y.-L. Lu, S. Mudanai, J. Park, A. Rahman, J. Rizk, W.-K. Shin, K. Soumyanath, H. Tashiro, C. Tsai, P. VanDerVoorn, J.-Y. Yeh, and P. Bai, “RF CMOS technology scaling in high- /metal gate era for RF SoC (system-onchip) applications,” in Proc. IEEE Int. Electron Device Meeting, Dec. 2010, pp. 27.2.1–27.2.4. [21] V. M. V. Serdio, Y. Azuma, S. Takeshita, T. Muraki, T. Teranishi, and Y. Majima, “Robust nanogap electrodes by self-terminating electroless gold plating,” Nanoscale, vol. 4, pp. 7161–7167, Sep. 2012. [22] V. M. V. Serdio, T. Muraki, S. Takeshita, D. H. S. Hurtado, S. Kano, T. Teranishi, and Y. Majima, “Gap separation-controlled nanogap electrodes by molecular ruler electroless gold plating,” RSC Advances, vol. 5, pp. 22160–22167, Feb. 2015. [23] H. Lee, Y. Liu, R. M. Westervelt, and D. Ham, “IC/microfluidic hybrid system for magnetic manipulation of biological cells,” IEEE J. SolidState Circuits, vol. 41, no. 6, pp. 1471–1480, Jun. 2006. [24] D. Kim, B. Goldstein, W. Tang, F. Sigworth, and E. Culurciello, “Noise analysis and performance comparison of low current measurement systems for biomedical applications,” IEEE Trans. Biomed. Circuits Syst., vol. 7, no. 1, pp. 52–62, Feb. 2013. [25] K. Gamo, K. Niitsu, and K. Nakazato, “A CMOS current integrator with 1.2 2.05 electroless-plated 1024 1024 microelectrode array for high-sensitivity bacteria detection,” in Proc. Int. Conf. Molecular Electronics and Bioelectronics, Jun. 2015. [26] J. Hasegawa, S. Uno, and K. Nakazato, “Amperometric electrochemical sensor array for on-chip simultaneous imaging: Circuit and microelectrode design consderations,” Jpn. J. Appl. Phys., vol. 50, no. 04DL03, p. 9, Apr. 2011. [27] T. Kuno, K. Niitsu, and K. Nakazato, “Amperometric electrochemical sensor array for on-chip simultaneous imaging,” Jpn. J. Appl. Phys., vol. 53, no. 04EL01, p. 7, Feb. 2014. [28] G. Zieve and S. Penman, “Small RNA species of the HeLa cell: Metabolism and subcellular localization,” Cell, vol. 8, no. 1, pp. 19–31, May 1976. [29] W. Ma, R. Zeng, D. Liu, Y. Zohar, and Y.-K. Lee, “SOI technologybased microfiltration system for circulating tumor cells isolation and enumeration,” in Proc. IEEE Int. Conf. Nano/Micro Engineered and Molecular Systems, Apr. 2014, pp. 333–336. [30] S. Ota, K. Niitsu, H. Kondo, M. Hori, and K. Nakazato, “A CMOS sensor platform with 1.2 2.05 electroless-plated 1024 1024 microelectrode array for high-sensitivity rapid direct bacteria counting,” in Proc. IEEE Biomedical Circuits and Systems Conf., Oct. 2014, pp. 460–463.
Kiichi Niitsu (S’06–M’10) was born in Japan, in 1983. He received the B.S. (summa cum laude), M.S. and Ph.D. degrees in electrical engineering from Keio University, Yokohama, Japan, in 2006, 2008, and 2010, respectively. From 2010, he was an Assistant Professor at Gunma University, Kiryu, Japan. Since 2012, he has been a Lecturer at Nagoya University, Nagoya, Japan. Since 2015, he has served concurrently as Precursory Research for Embryonic Science and Technology (PRESTO) Researcher, Japan Science and Technology Agency (JST), Tokyo, Japan. His current research interest lies in the low-power and high-speed technologies of analog and digital VLSI circuits for biomedical application. From 2008 to 2010, he was a Research Fellow of the Japan Society for the Promotion of Science (JSPS), a Research Assistant in the Global Center of Excellence (GCOE) Program, Keio University, and a Collaboration Researcher in the Keio Advanced Research Center (KARC). Dr. Niitsu was awarded the 2006 KEIO KOUGAKUKAI Award, the 2007 INOSE Science Promotion Award, the 2008 IEEE SSCS Japan Chapter Young Researcher Award, and the 2009 IEEE SSCS Japan Chapter Academic Research Award, both from the IEEE Solid-State Circuits Society Japan Chapter. He also received the 2008 FUJIWARA Award from the FUJIWARA Foundation, 2011 YASUJIRO NIWA Outstanding Paper Award, 2011 FUNAI Research Promotion Award, 2011 Ando Incentive Prize for the Study of Electronics, 2011 Ericsson Young Scientist Award, 2012 ASP-DAC University LSI Design Contest Design Award, NF Foundation R&D Encouragement Award, AKASAKI Award from Nagoya University, and IEEE Nagoya Section Young Researcher Award. He served as a Review Committee Member of APCCAS 2014, and on the editorial committee of the Institute of Electronics, Information and Communication Engineers of Japan (IEICE) Transactions on Electronics, Special Section on Analog Circuits and Related SoC Integration Technologies, and IEICE ESS Fundamental Review. He is a member of the IEICE and Japan Society of Applied Physics (JSAP).
Shoko Ota received the B.S. and M.S. degrees in electrical engineering and computer science from Nagoya University, Nagoya, Japan, in 2013 and 2015, respectively. Currently, she is with Brother Industries, Ltd., Nagoya, Japan. Her research activity was focused on mixed-signal CMOS integrated circuits for biomedical applications.
Kohei Gamo (S’15) was born in Osaka, Japan, in 1993. He received the B.S. degree in electrical engineering and computer science from Nagoya University, Nagoya, Japan, in 2015. Currently, he is working toward the M.S. degree at Nagoya University. His research activity is focused on mixed-signal CMOS integrated circuits for biomedical applications.
Hiroki Kondo was born in Okazaki, Japan, in 1971. He received the B.E., M.E., and Ph.D. degrees in material science from Nagoya University, Nagoya, Japan, in 1994, 1996, and 1999, respectively. He joined Fujitsu Corporation, starting research on reliability technology for multilayer interconnects in ULSI. He returned to Nagoya University as an Assistant Professor and began research on high-k/metal gate stack technologies. In particular, he focused on plasma nitridation processes. He has also studied formation processes of high-k gate stack structures using scanning tunneling microscopy (STM), transmission electron microscopy (TEM), and synchrotron hard-X-ray photoelectron spectroscopy.
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In 2009, he joined the Plasma Nanotechnology Research Center, Nagoya University, researching high-density radical source and its application to nitride semiconductor growth. He also studies advanced plasma processes for nanomaterials and nanodevices, especially carbon nanomaterials and their applications. He had 37 invited talks in international conferences and authored more than 110 international academic journal papers.
Masaru Hori was born in Gifu, Japan, in 1958. He received the B.E. and M. Tech degrees from the Department of Electric Engineering at Waseda University, Tokyo, Japan, in 1981 and 1983, respectively, and the Ph.D. degree from the Department of Electric Engineering at Nagoya University, Nagoya, Japan, in 1986. In 1986, he joined Toshiba Corporation, Kawasaki, Japan. Since 1992, he has been with the Graduate School of Engineering, Nagoya University. He became a Research Associate in 1992, an Assistant Professor in 1994, an Associate Professor in 1996, a Researcher at Cavendish Laboratory, University of Cambridge, Cambridge, U.K., in 1997, and a Professor in 2004. He was Vice Director at the Plasma Nanotechnology Research Centre, Nagoya University, in 2007, Director, Plasma Nanotechnology Research Centre, Nagoya University, in 2009, and Director of Plasma Medical Science Global Innovation Centre, Nagoya University, in 2013. His current research interests include plasma electronics, plasma processing science in materials and devices, and plasma life sciences to medicine and agriculture.
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Dr. Hori is a Fellow of The Japan Society of Applied Physics. He was the recipient of the Prize for Science and Technology by the Minister of Education, Culture, Sports, Science and Technology, Japan, in 2011, Prize for IndustryGovernment-Academia Cooperation Person of Merit, Science and Technology Minister Policy Award, Japan, in 2011, and the 14th Plasma Material Science Award, Division of 153 Committee, Japan Society for the Promotion of Science, Japan, in 2012.
Kazuo Nakazato was born on October 18, 1952, in Japan. He received the B.S., M.S., and Ph.D. degrees in physics from the University of Tokyo, Tokyo, Japan, in 1975, 1977, and 1980, respectively. In 1981, he joined the Central Research Laboratory, Hitachi Ltd., Tokyo, Japan, working on high-speed silicon self-aligned bipolar devices [sidewall base contact structure (SICOS)], which were adopted in main frame computer Hitachi M-880/420. In 1989, he moved to Hitachi Cambridge Laboratory, Hitachi Europe Ltd., Cambridge, U.K., as a Senior Researcher and a Laboratory Manager, working on experimental and theoretical study of quantum electron transport in semiconductor nano structures, including single-electron memory. Since 2004, he has been a Professor of Intelligent Device in the Department of Electrical Engineering and Computer Science, Graduate School of Engineering, Nagoya University, Nagoya, Japan. His main interests are BioCMOS technology, single molecule-CMOS hybrid devices, and CMOS analog circuits for integrated sensors.
