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Particle concentration sensor using control volume between double electrical sensing zones
Current Applied Physics 6S1 (2006) e232–e236 www.elsevier.com/locate/cap www.kps.or.kr
Particle concentration sensor using control volume between double electrical sensing zones D.W. Lee, S. Yi, Y.-H. Cho
*
Digital Nanolocomotion Center, Department of BioSystems, Daejeon 305 701, Republic of Korea Received 15 July 2005 Available online 18 April 2006
Abstract We present a novel flow-rate independent cell concentration sensor using a fixed control volume between double electrical sensing zones. Compared to Hemacytometer [L.M. Prescott, J.P. Harley, D.A. Klein, Microbiology, McGraw-Hill, New York, 2002.] and Coulter Counter [M. Koch, A.G.R. Evans, A. Brunnschweiler, Design and fabrication of a micromachined coulter counter, J. Micromech. Microeng. 9 (1999) 159–161] requiring an accurate fluid volume measurement or precision flow-rate control, the present cell concentration sensor offers the flow-rate independent method for the cell concentration measurement with counting cells in a fixed control volume of 0.423 ll. In the experimental study, we have used the RBC samples of three different concentrations and compared the results obtained from the present device with those from Hemacytometer. Using the fabricated devices, we have made two different cell concentration measurements: (1) with single electrical sensing at a fixed flow-rate of 1.2 ll/min; (2) with double electrical sensing with a known control volume of 0.423 ± 0.01 ll. Compared to Hemacytometer, the single and double sensing methods show the maximum concentration errors of 8.7% and 10.3%, which are in the measurement error range of Hemacytometer. We also measure the cell concentration within the maximum concentration errors of 10.3% in two cases: (1) two different flow-rates of 5 and 0.5 ll/min, (2) the varying flow-rates from 2 to 1 ll/min, respectively. Therefore, we verify the flow-rate independent measurement capability of the present device. Finally, we conclude that the present sensor can measure cell concentration without the accurate control and measurement of flow-rate. 2006 Elsevier B.V. All rights reserved. PACS: 85.85.+j; 87.80.y; 87.80.Tq Keywords: Cell counter; Coulter counter; Electrical sensing zone; Cell concentration sensor
1. Introduction Cell concentration measurement is a basic process for medical diagnosis and cell research. Therefore, many cell concentration measurement devices have been developed at the commercial level. They may be divided broadly into two categories. First one, Hemacytometer [1] is a special slide glass that has a grid to count cells in a fixed volume of 0.1 ll. While it is portable and cheap, the process of cell counting is manual and requires an optical microscope.
*
Corresponding author. Fax: +82 42 869 8690. E-mail address: [email protected] (Y.-H. Cho).
1567-1739/$ - see front matter 2006 Elsevier B.V. All rights reserved. doi:10.1016/j.cap.2006.01.046
Compared to Hemacytometer, second one, Coulter Counters [2,3] can automatically count the cell passing through an electrical sensing zone. They measure cell concentration based on two factors: (1) cell number passing through the electrical sensing zone per unit time; (2) flow-rate. Since the cell concentration measurement with Coulter Counter is based on flow-rate or fluidic volume passing through the electrical sensing zone, the accuracy of the cell concentration measurement is dependent on flow-rate control. Thus, this method requires accurate fluidic control systems, which are expensive and require a large space. In this paper, we propose a novel cell concentration sensor composed of the fixed control volume and double electrical sensing zones as shown in Fig. 1. In order to measure
D.W. Lee et al. / Current Applied Physics 6S1 (2006) e232–e236
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Fig. 3. Principle of electrical sensing zone.
Fig. 1. The schematic view of the present cell concentration sensing chip.
cell concentration without accurate fluidic control systems, the present cell concentration sensor counts the number of cells in the fixed control volume between double electrical sensing zones. Since the control volume is initially empty, the number of cell (Ncv = Nin Nout) in the fixed control volume increases initially but converges to a specific number as shown in Fig. 2. The cell concentration is easily calculated from the converged number of cell in the known control volume without accurate fluidic control systems. Because the number of cells in the fixed control volume does not change regardless of change in the flow-rate, this method is insensitive to flow-rate.
Fig. 2. Cell concentration measurement principle based on the total cell number of fixed control volume.
2. Cell counting principle The principle of cell counting relies on resistance change of the orifice channel in the electrical sensing zone, when particles pass through the orifice channel. Since the resistance change, DR, causes of polarization effects at the particle/electrolyte interface preventing any current flow through the particle itself, it depends on the integral of the particle cross-section as shown in Fig. 3. The resistance change of the element d (DR) over a finite distance dl can be determined [4] by q a dl . a dðDRÞ ¼ f 2 1 ; ð1Þ A A where qf, A and a are the electrical resistivity of the fluid, the cross-sectional area of the orifice channel, and the cross-sectional area of the particle, respectively. For a diameter of spherical particle, d, and a diameter of orifice channel, D, the resistance change for an element of thickness, dl, at a distance, l, from the center of the sphere may be determined and this can be integrated to give the resistance change due to the particle [4], as shown below 2 3 Z 2l 1 4q 6 sin ðd=DÞ d7 ð2Þ DR ¼ 2 dðDRÞ ¼ f 4qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi 5. D pD 2 0 1 ðd=DÞ From Eq. (2), when a particle passes through the orifice channel in the electrical sensing zone, an electrical resistance of the orifice channel changes and creates a pulse of voltage by electrodes integrated with microchannel, as
Fig. 4. Electrical sensing of the cell: (a) electrical sensing zone (Coulter Counter) and (b) detected signals.
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shown in Fig.4(a). The amplitudes of voltage in Fig. 4(b) are not same, due to the different volume of cells, but higher than the threshold voltage. One pulse of which amplitude is higher than the threshold voltage means that one cell pass the orifice channel and the number of cell passing the orifice channel is calculated by counting all pulse of which amplitude is larger than the threshold voltage. 3. Design and fabrication process 3.1. Design The present cell concentration sensor is composed of the control volume and double electrical sensing zones before and after the control volume. Both control volume and double electrical sensing zones made of PDMS microchanel are on a single chip. The control volume is 0.423 ll, which is similar to the volume of Hemacytometer. Based on the volume of a RBC (red blood cell) sample, 60– 120 fl, the microchannel of electrical sensing zone has a height of 10 lm, a width of 10 lm, and a length of 40 lm and sensing electrode for measuring resistance of microchannel has a thickness of 0.2 lm, a width of 5 lm as shown in Fig. 4(a). From these channel dimensions and buffer solution (phosphate buffered saline: conductivity of 1.6 S/m), we can estimate a channel resistance of 0.25 MX, a resistance change of 5.9 (60 fl)–13.8 (120 fl) kX, and a voltage change of 1.8 (60 fl)–4.21 (120 fl) V according to the passing RBC. 3.2. Fabrication process The present cell concentration sensor, consisting of the control volume and double electrical sensing zones, has been fabricated by three steps: (1) fabrication of electrodes on the glass substrate; (2) fabrication of a PDMS (Polydimethylsiloxane) microchannel; (3) bonding of them. Fig. 5 shows the fabrication process of the present cell concentration sensor. For the fabrication of electrodes on the glass substrates, we use 400 Pyrex glass wafers. We deposit the sputtered Cr and Au layer on the glass wafers for elec˚ and 1500 A ˚ , respectively trodes at the thickness of 500 A and then pattern. The PDMS microchannel is fabricated by micromolding technique. The mold is fabricated on
Fig. 6. The fabricated device.
Fig. 7. Experimental apparatus for the present cell concentration sensor using the control volume between double electrical sensing zones.
the silicon substrate by using 10 lm-thick positive PR, AZ9260. Finally, we bond a PDMS microchannel and the glass substrate using high frequency generator, BD10AS (Electro-Technic Products, INC). Fig. 6 shows the fabricated cell concentration sensor and enlarged views with measured dimensions. The experimental apparatus has been designed (Fig. 7) for the evaluation of the fabricated cell concentration sensor. The bias voltage is applied between inlet and outlet for measuring resistance change as voltage change. Signals from electrical sensing zones are amplified with differential amplifier and enter A/D board through the filter. We use A/D board of PCI-6251 (National Instrument, INC) to acquire data. 4. Experimental results and discussion The experimentally measured signals, when RBC passes through electrical sensing zone, are compared to the theoretically estimated signals in the PBS (phosphate buffered saline) medium of 1.6 S/m conductivity as shown in Table 1. The signal variation of a measured signal, as shown in
Table 1 Electrical property of electrical sensing zone at buffer (PBS) Conductivity of 1.6 S/m
Fig. 5. Fabrication process of the electrical sensing zones.
Theoretical Experimental
R (kX)
DR (kX)
DV (V)
250 –
5.9 (60 fl)–13.8 (120 fl) –
3 (60 fl)–7 (120 fl) 1.5–5
D.W. Lee et al. / Current Applied Physics 6S1 (2006) e232–e236
Fig. 8. Measured electrical signals from single electrical sensing zone.
Table 2 Comparison of the experimental results SESZa
Hemacytometer 5
5.80 ± 0.7 [ · 10 cells/ml] 7.70 ± 0.7 [ · 105 cells/ml] 1.15 ± 0.17 [ · 106 cells/ml]
DESZb c
5.67 ± 0.30 (2.2%) 7.30 ± 0.46 (5.2%) 1.25 ± 0.12 (8.7%)
5.20 ± 0.15 (10.3%) 7.49 ± 1.16 (2.7%) 1.21 ± 0.05 (5.2%)
a
SESZ: single electrical sensing zone at flow-rate of 1.2 ll/min. DESZ: double electrical sensing zone at control volume of 22.9 ± 0.98 ll. c Maximum error of mean value compared to Hemacytometer. b
Fig. 8, results from the RBC volume distribution. Although there is noise around 1 (V), the signal can be distinguished from the noise. The minimum voltage measured
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from electrical sensing zone is about 1.8 V, thus we can determine the threshold voltage of 1.5 (V) to count RBC. We have used the RBC samples of three different concentrations as shown in Table 2 and compared the results measured from the fabricated device with those from Hemacytometer. To test the fabricated devices, we have measured cell concentration in two ways: (1) Cell concentration measurement with single sensing method which uses single electrical sensing zone at a fixed flow-rate of 1.2 ll/ min; (2) Cell concentration measurement using double electrical sensing zones with the fixed control volume of 0.423 ± 0.01 ll. The results obtained from the single sensing method and the double sensing method is shown in Fig. 9. Table 2 compares the obtained cell concentration from two different devices with the cell concentration of Hemacytometer for three different RBC samples. Compared to Hemacytometer, the single and double sensing method show the maximum errors of 8.7 and 10.3% as shown in Table 2, which are in the measurement error range of Hemacytometer. In order to verify if the present device is the flow-rate independent (or insensitive) at cell concentration measurement, we measure the cell concentration using the present device under the various flow-rate conditions: (1) the different flow-rate of 5 and 0.5 ll/min; (2) the flow-rate change
Fig. 9. Measured cell number from fabricated devices: (a) single sensing method using single electrical sensing zone and flow-rate of 1.2 ll/min and (b) double sensing method using double electrical sensing zones and control volume of 0.423 ± 0.01 ll.
Fig. 10. Measured cell number from cell concentration sensor using double electrical sensing zones and the control volume of 0.423 ± 0.01 ll: (a) at the constant different flow-rates and (b) at the change of flow-rate.
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from 2 ll/min to 1 ll/min during measuring cell concentration. Fig. 10 shows the results obtained under the various flow-rate conditions, which is within the maximum errors of 10.3%. From these tests, we verify that the performance of the present device is not affected by the fluidic control. 5. Conclusions We present a novel cell concentration sensor using a fixed control volume between double electrical sensing zones. The previous methods, Hemacytometer [1] and Coulter Counter [2], require an accurate fluid volume measurement or precision flow-rate control. The present cell concentration sensor, however, offers a simple and automated method for the cell concentration measurement without an accurate fluid measurement and precision flow-rate control. We have used the RBC samples of three different concentrations and compared the results obtained from the present device with those from Hemacytometer. Using the fabricated devices, we have made two different cell concentration measurements: (1) using single electrical sensing at a fixed flow-rate of 1.2 ll/min; (2) using double electrical sensing with a known control volume of 0.423 ± 0.01 ll. Compared to Hemacytometer, the single and double sensing methods show the maximum concentration errors of 8.7% and 10.3%, which are in the measurement error range of Hemacytometer. We also measure the cell concentration within the maximum concentration
errors of 10.3% in two cases: (1) two different flow-rates of 5 and 0.5 ll/min; (2) the varying flow-rates from 2 to 1 ll/min, respectively. Therefore, we demonstrate that the present cell concentration sensor results in a simple and automated method for the cell concentration measurement without an accurate fluid measurement and precision flowrate control. The present cell concentration sensor does not require flow sensors or accurate pumps, thus increasing the simplicity and adaptability for an integrated system while reducing the system size and cost. Acknowledgements This work has been supported by the National Creative Research Initiative Program of the Ministry of Science and Technology (MOST) under the project title of ‘‘Realization of Bio-Inspired Digital Nanoactuators’’. References [1] L.M. Prescott, J.P. Harley, D.A. Klein, Microbiology, McGraw-Hill, New York, 2002. [2] M. Koch, A.G.R. Evans, A. Brunnschweiler, Design and Fabrication of a Micromachined Coulter Counter, J. Micromech. Microeng. 9 (1999) 159–161. [3] U.D. Larsen, G. Blankenstein, J. Branebjerg, Microchip Coulter Particle Counter, in: Proceedings of International Solid State Sensors and Actuators Conference 1997, Chicago, USA, (1997), pp. 1319–1322. [4] T. Allen, Particles Size Measurements, Chapman and Hall, New York, 1990.
Particle concentration sensor using control volume between double electrical sensing zones D.W. Lee, S. Yi, Y.-H. Cho
*
Digital Nanolocomotion Center, Department of BioSystems, Daejeon 305 701, Republic of Korea Received 15 July 2005 Available online 18 April 2006
Abstract We present a novel flow-rate independent cell concentration sensor using a fixed control volume between double electrical sensing zones. Compared to Hemacytometer [L.M. Prescott, J.P. Harley, D.A. Klein, Microbiology, McGraw-Hill, New York, 2002.] and Coulter Counter [M. Koch, A.G.R. Evans, A. Brunnschweiler, Design and fabrication of a micromachined coulter counter, J. Micromech. Microeng. 9 (1999) 159–161] requiring an accurate fluid volume measurement or precision flow-rate control, the present cell concentration sensor offers the flow-rate independent method for the cell concentration measurement with counting cells in a fixed control volume of 0.423 ll. In the experimental study, we have used the RBC samples of three different concentrations and compared the results obtained from the present device with those from Hemacytometer. Using the fabricated devices, we have made two different cell concentration measurements: (1) with single electrical sensing at a fixed flow-rate of 1.2 ll/min; (2) with double electrical sensing with a known control volume of 0.423 ± 0.01 ll. Compared to Hemacytometer, the single and double sensing methods show the maximum concentration errors of 8.7% and 10.3%, which are in the measurement error range of Hemacytometer. We also measure the cell concentration within the maximum concentration errors of 10.3% in two cases: (1) two different flow-rates of 5 and 0.5 ll/min, (2) the varying flow-rates from 2 to 1 ll/min, respectively. Therefore, we verify the flow-rate independent measurement capability of the present device. Finally, we conclude that the present sensor can measure cell concentration without the accurate control and measurement of flow-rate. 2006 Elsevier B.V. All rights reserved. PACS: 85.85.+j; 87.80.y; 87.80.Tq Keywords: Cell counter; Coulter counter; Electrical sensing zone; Cell concentration sensor
1. Introduction Cell concentration measurement is a basic process for medical diagnosis and cell research. Therefore, many cell concentration measurement devices have been developed at the commercial level. They may be divided broadly into two categories. First one, Hemacytometer [1] is a special slide glass that has a grid to count cells in a fixed volume of 0.1 ll. While it is portable and cheap, the process of cell counting is manual and requires an optical microscope.
*
Corresponding author. Fax: +82 42 869 8690. E-mail address: [email protected] (Y.-H. Cho).
1567-1739/$ - see front matter 2006 Elsevier B.V. All rights reserved. doi:10.1016/j.cap.2006.01.046
Compared to Hemacytometer, second one, Coulter Counters [2,3] can automatically count the cell passing through an electrical sensing zone. They measure cell concentration based on two factors: (1) cell number passing through the electrical sensing zone per unit time; (2) flow-rate. Since the cell concentration measurement with Coulter Counter is based on flow-rate or fluidic volume passing through the electrical sensing zone, the accuracy of the cell concentration measurement is dependent on flow-rate control. Thus, this method requires accurate fluidic control systems, which are expensive and require a large space. In this paper, we propose a novel cell concentration sensor composed of the fixed control volume and double electrical sensing zones as shown in Fig. 1. In order to measure
D.W. Lee et al. / Current Applied Physics 6S1 (2006) e232–e236
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Fig. 3. Principle of electrical sensing zone.
Fig. 1. The schematic view of the present cell concentration sensing chip.
cell concentration without accurate fluidic control systems, the present cell concentration sensor counts the number of cells in the fixed control volume between double electrical sensing zones. Since the control volume is initially empty, the number of cell (Ncv = Nin Nout) in the fixed control volume increases initially but converges to a specific number as shown in Fig. 2. The cell concentration is easily calculated from the converged number of cell in the known control volume without accurate fluidic control systems. Because the number of cells in the fixed control volume does not change regardless of change in the flow-rate, this method is insensitive to flow-rate.
Fig. 2. Cell concentration measurement principle based on the total cell number of fixed control volume.
2. Cell counting principle The principle of cell counting relies on resistance change of the orifice channel in the electrical sensing zone, when particles pass through the orifice channel. Since the resistance change, DR, causes of polarization effects at the particle/electrolyte interface preventing any current flow through the particle itself, it depends on the integral of the particle cross-section as shown in Fig. 3. The resistance change of the element d (DR) over a finite distance dl can be determined [4] by q a dl . a dðDRÞ ¼ f 2 1 ; ð1Þ A A where qf, A and a are the electrical resistivity of the fluid, the cross-sectional area of the orifice channel, and the cross-sectional area of the particle, respectively. For a diameter of spherical particle, d, and a diameter of orifice channel, D, the resistance change for an element of thickness, dl, at a distance, l, from the center of the sphere may be determined and this can be integrated to give the resistance change due to the particle [4], as shown below 2 3 Z 2l 1 4q 6 sin ðd=DÞ d7 ð2Þ DR ¼ 2 dðDRÞ ¼ f 4qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi 5. D pD 2 0 1 ðd=DÞ From Eq. (2), when a particle passes through the orifice channel in the electrical sensing zone, an electrical resistance of the orifice channel changes and creates a pulse of voltage by electrodes integrated with microchannel, as
Fig. 4. Electrical sensing of the cell: (a) electrical sensing zone (Coulter Counter) and (b) detected signals.
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D.W. Lee et al. / Current Applied Physics 6S1 (2006) e232–e236
shown in Fig.4(a). The amplitudes of voltage in Fig. 4(b) are not same, due to the different volume of cells, but higher than the threshold voltage. One pulse of which amplitude is higher than the threshold voltage means that one cell pass the orifice channel and the number of cell passing the orifice channel is calculated by counting all pulse of which amplitude is larger than the threshold voltage. 3. Design and fabrication process 3.1. Design The present cell concentration sensor is composed of the control volume and double electrical sensing zones before and after the control volume. Both control volume and double electrical sensing zones made of PDMS microchanel are on a single chip. The control volume is 0.423 ll, which is similar to the volume of Hemacytometer. Based on the volume of a RBC (red blood cell) sample, 60– 120 fl, the microchannel of electrical sensing zone has a height of 10 lm, a width of 10 lm, and a length of 40 lm and sensing electrode for measuring resistance of microchannel has a thickness of 0.2 lm, a width of 5 lm as shown in Fig. 4(a). From these channel dimensions and buffer solution (phosphate buffered saline: conductivity of 1.6 S/m), we can estimate a channel resistance of 0.25 MX, a resistance change of 5.9 (60 fl)–13.8 (120 fl) kX, and a voltage change of 1.8 (60 fl)–4.21 (120 fl) V according to the passing RBC. 3.2. Fabrication process The present cell concentration sensor, consisting of the control volume and double electrical sensing zones, has been fabricated by three steps: (1) fabrication of electrodes on the glass substrate; (2) fabrication of a PDMS (Polydimethylsiloxane) microchannel; (3) bonding of them. Fig. 5 shows the fabrication process of the present cell concentration sensor. For the fabrication of electrodes on the glass substrates, we use 400 Pyrex glass wafers. We deposit the sputtered Cr and Au layer on the glass wafers for elec˚ and 1500 A ˚ , respectively trodes at the thickness of 500 A and then pattern. The PDMS microchannel is fabricated by micromolding technique. The mold is fabricated on
Fig. 6. The fabricated device.
Fig. 7. Experimental apparatus for the present cell concentration sensor using the control volume between double electrical sensing zones.
the silicon substrate by using 10 lm-thick positive PR, AZ9260. Finally, we bond a PDMS microchannel and the glass substrate using high frequency generator, BD10AS (Electro-Technic Products, INC). Fig. 6 shows the fabricated cell concentration sensor and enlarged views with measured dimensions. The experimental apparatus has been designed (Fig. 7) for the evaluation of the fabricated cell concentration sensor. The bias voltage is applied between inlet and outlet for measuring resistance change as voltage change. Signals from electrical sensing zones are amplified with differential amplifier and enter A/D board through the filter. We use A/D board of PCI-6251 (National Instrument, INC) to acquire data. 4. Experimental results and discussion The experimentally measured signals, when RBC passes through electrical sensing zone, are compared to the theoretically estimated signals in the PBS (phosphate buffered saline) medium of 1.6 S/m conductivity as shown in Table 1. The signal variation of a measured signal, as shown in
Table 1 Electrical property of electrical sensing zone at buffer (PBS) Conductivity of 1.6 S/m
Fig. 5. Fabrication process of the electrical sensing zones.
Theoretical Experimental
R (kX)
DR (kX)
DV (V)
250 –
5.9 (60 fl)–13.8 (120 fl) –
3 (60 fl)–7 (120 fl) 1.5–5
D.W. Lee et al. / Current Applied Physics 6S1 (2006) e232–e236
Fig. 8. Measured electrical signals from single electrical sensing zone.
Table 2 Comparison of the experimental results SESZa
Hemacytometer 5
5.80 ± 0.7 [ · 10 cells/ml] 7.70 ± 0.7 [ · 105 cells/ml] 1.15 ± 0.17 [ · 106 cells/ml]
DESZb c
5.67 ± 0.30 (2.2%) 7.30 ± 0.46 (5.2%) 1.25 ± 0.12 (8.7%)
5.20 ± 0.15 (10.3%) 7.49 ± 1.16 (2.7%) 1.21 ± 0.05 (5.2%)
a
SESZ: single electrical sensing zone at flow-rate of 1.2 ll/min. DESZ: double electrical sensing zone at control volume of 22.9 ± 0.98 ll. c Maximum error of mean value compared to Hemacytometer. b
Fig. 8, results from the RBC volume distribution. Although there is noise around 1 (V), the signal can be distinguished from the noise. The minimum voltage measured
e235
from electrical sensing zone is about 1.8 V, thus we can determine the threshold voltage of 1.5 (V) to count RBC. We have used the RBC samples of three different concentrations as shown in Table 2 and compared the results measured from the fabricated device with those from Hemacytometer. To test the fabricated devices, we have measured cell concentration in two ways: (1) Cell concentration measurement with single sensing method which uses single electrical sensing zone at a fixed flow-rate of 1.2 ll/ min; (2) Cell concentration measurement using double electrical sensing zones with the fixed control volume of 0.423 ± 0.01 ll. The results obtained from the single sensing method and the double sensing method is shown in Fig. 9. Table 2 compares the obtained cell concentration from two different devices with the cell concentration of Hemacytometer for three different RBC samples. Compared to Hemacytometer, the single and double sensing method show the maximum errors of 8.7 and 10.3% as shown in Table 2, which are in the measurement error range of Hemacytometer. In order to verify if the present device is the flow-rate independent (or insensitive) at cell concentration measurement, we measure the cell concentration using the present device under the various flow-rate conditions: (1) the different flow-rate of 5 and 0.5 ll/min; (2) the flow-rate change
Fig. 9. Measured cell number from fabricated devices: (a) single sensing method using single electrical sensing zone and flow-rate of 1.2 ll/min and (b) double sensing method using double electrical sensing zones and control volume of 0.423 ± 0.01 ll.
Fig. 10. Measured cell number from cell concentration sensor using double electrical sensing zones and the control volume of 0.423 ± 0.01 ll: (a) at the constant different flow-rates and (b) at the change of flow-rate.
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from 2 ll/min to 1 ll/min during measuring cell concentration. Fig. 10 shows the results obtained under the various flow-rate conditions, which is within the maximum errors of 10.3%. From these tests, we verify that the performance of the present device is not affected by the fluidic control. 5. Conclusions We present a novel cell concentration sensor using a fixed control volume between double electrical sensing zones. The previous methods, Hemacytometer [1] and Coulter Counter [2], require an accurate fluid volume measurement or precision flow-rate control. The present cell concentration sensor, however, offers a simple and automated method for the cell concentration measurement without an accurate fluid measurement and precision flow-rate control. We have used the RBC samples of three different concentrations and compared the results obtained from the present device with those from Hemacytometer. Using the fabricated devices, we have made two different cell concentration measurements: (1) using single electrical sensing at a fixed flow-rate of 1.2 ll/min; (2) using double electrical sensing with a known control volume of 0.423 ± 0.01 ll. Compared to Hemacytometer, the single and double sensing methods show the maximum concentration errors of 8.7% and 10.3%, which are in the measurement error range of Hemacytometer. We also measure the cell concentration within the maximum concentration
errors of 10.3% in two cases: (1) two different flow-rates of 5 and 0.5 ll/min; (2) the varying flow-rates from 2 to 1 ll/min, respectively. Therefore, we demonstrate that the present cell concentration sensor results in a simple and automated method for the cell concentration measurement without an accurate fluid measurement and precision flowrate control. The present cell concentration sensor does not require flow sensors or accurate pumps, thus increasing the simplicity and adaptability for an integrated system while reducing the system size and cost. Acknowledgements This work has been supported by the National Creative Research Initiative Program of the Ministry of Science and Technology (MOST) under the project title of ‘‘Realization of Bio-Inspired Digital Nanoactuators’’. References [1] L.M. Prescott, J.P. Harley, D.A. Klein, Microbiology, McGraw-Hill, New York, 2002. [2] M. Koch, A.G.R. Evans, A. Brunnschweiler, Design and Fabrication of a Micromachined Coulter Counter, J. Micromech. Microeng. 9 (1999) 159–161. [3] U.D. Larsen, G. Blankenstein, J. Branebjerg, Microchip Coulter Particle Counter, in: Proceedings of International Solid State Sensors and Actuators Conference 1997, Chicago, USA, (1997), pp. 1319–1322. [4] T. Allen, Particles Size Measurements, Chapman and Hall, New York, 1990.