1. Introduction

In order to count and categorize heterogeneous cell populations and to measure the physical characteristics of a single cell, flow cytometry is widely used in microbiological detection and medical diagnosis^[1]. In particular, the number of different circulating cells (red blood cells (RBCs), white blood cells (WBCs), platelets, etc.) in the blood is crucial for disease prevention, diagnosis, and treatment^[2]. At present, the mainstream solution for cell counting is optical flow cytometry, which can provide a wide range of information about cells. However, complicated optics make it difficult to meet the needs of point-of-care testing (POCT) and diagnosis, due to its high cost and bulky device size^{[3, 4]}. In contrast to the optical approach, the electrical Coulter-counters shows great advantages of low cost and miniaturization. However, it can only provide limited information about cell size and density^{[5, 6]}. To address the increasing requirements for POC diagnosis, it is imperative to create a high-throughput label-free flow cytometer with single-cell resolution.

Electrical impedance spectroscopy (EIS) is frequently used to describe the dielectric characteristics of mediums such as fuel cells^[7], lithium-ion batteries^[8], and biological systems^[9−12]. Several attempts have been made to combine EIS with microfluidics for the characterization of a single cell^[13−16]. The impedance variations of cells flowing through the induction electrode in the microchannel are converted into current signals by applying a pre-set frequency AC excitation voltage, which is subsequently amplified and demodulated by the circuit^[17]. The external syringe pump was employed to generate liquid flow in impedance flow cytometry investigations in the past^{[18, 19]}. However, making multi-channel integration using such equipment is impossible. In our previous work^[20], we achieved a high-throughput cell-counting chip with 100 cell counting channels using a thermal-bubble micropump as a driving source for the first time. Nevertheless, much as in the majority of previous research, we extracted and processed the signal using an impedance analyser, which significantly worsened the detection limits and integration^{[21, 22]}. In contrast, complementary metal-oxide−semiconductor (CMOS) silicon integrated circuits that has been proved to have smaller size and lower parasitic noise for multichannel deployment can be used to measure EIS^{[23, 24]}.

Here, based on integrated CMOS-micro-electromechanical systems (MEMS) process, we developed an impedance flow cytometer, which includes a thermal bubble micropump, parallel three electrodes, microchannel, TIA on-chip, and auxiliary circuit. In order to monitor the solutions with different impedances, we designed trans-impedance amplifier (TIA) to achieve 5−60 kΩ trans-impedance adjustable. At the same time, the input end of the on-chip TIA is directly connected to the electrodes, and the CMOS-MEMS single-chip architecture eliminates the parasitic parameters generated on the interface between MEMS and CMOS in the traditional sensor. In this way, the signal-to-noise ratio (SNR) can be improved effectively. Various sizes polystyrene beads were used to test the counting accuracy and speed to assess the performance of the chip. Meanwhile, the function of the chip to classify polystyrene beads of various diameters was also assessed. We also used MSTO-211H cells to check the accuracy of the cell-counting performance of the proposed chip.

2. Experiment

2.1 EIS flow cytometry fundamentals

Differential EIS systems can be used to detect impedance changes in liquids, and have greater SNR and detection limits than single-ended EIS^[25]. As shown in Fig. 1, the middle electrode (ME) receives the excitation AC voltage signal, whereas the top electrode (TE) and bottom electrode (BE) measure the induced current signal. The dielectric characteristics of the cell can be equivalent to the parallel connection of the capacitor formed by the membrane (phospholipid bilayer) and the resistor formed by the cytoplasm in a specific frequency range (0.5−5 MHz)^[17]. In the microfluidic channel, the flow of cells through the induction electrode causes the difference between the impedance of TE to ME and BE to ME, resulting in a change in the induction current. The changing current value $\Delta I$ can be calculated as Eq. (1):

$$\Delta I=\frac{V}{{Z}_{0}+{Z}_\rm{cell}}-\frac{V}{{Z}_{0}}.$$

(1)

Here Z₀ represents the equivalent impedance of the cell suspension plus the capacity of the double-layer capacitance (C_DL). C_DL is formed between the metal electrode and the cell suspension. Z₀ can be calculated from Eq. (2):

$${Z}_{0}={R}_\rm{solution}+2\frac{1}{jw\left({C}_\rm{DL}\right)}.$$

(2)

As can be seen from Fig. 1, Z_cell consists of membrane capacitor (C_membrane) and cytoplasm resistor (R_cytoplasm), the formula should be

$${Z}_\rm{cell}={R}_\rm{cytoplasm}+\frac{1}{jw\left({C}_\rm{membrane}\right)}.$$

(3)

As a result, the impedance change can be inferred in the flow channel when the cell passes by examining the differential current values of the electrode plates on both sides.

2.2 EIS flow cytometry design

Fig. 2(a) presents the basic structure of the microfluidic part of the chip, which mainly includes a micro-flow channel, micro-pillar array, coplanar three electrodes, and thermal bubble micropump. The width and height of the micro-flow channel are adjusted to 30 and 30 μm, respectively, making it slightly larger than the size of the cell to facilitate the smooth passage of single cell and prevent the error caused by the overlapping passage of multi-cells. The coplanar electrode in the production process has less alignment difficulty than the opposite electrode of the classic Coulter-counter^[19], which can effectively increase detection accuracy and decrease the cost. Additionally, the three-electrode differential detection mode can increase the detection range by eliminating common mode interference in comparison to the double electrodes. Tantalum was selected as the electrode material in this work because of its excellent chemical stability, which is not easily corroded by cell suspension, and its compatibility with standard CMOS-MEMS processes^[26]. The size of the electrode and spacing of the electrodes affect the value of the reference impedance Z₀, and the smaller the Z₀, the more sensitive the electrode response is to the cell impedance Z_cell. Z₀ decreases as the electrode width increases and is directly proportional to the distance between electrodes. This relationship has been consistently demonstrated in previous research^[20]. To investigate the connection between electrode size and response amplitude, three sets of electrodes with widths and spacings of 30/15, 30/10, and 25/15 μm were employed in our previous experiment. Consistent with the theoretical derivation, coplanar electrodes with a 10 μm spacing and a width of 30 μm exhibited the highest response. Therefore, we set the electrode length as 30 μm and the space between two electrodes as 10 μm. Rather than using the conventional external peristaltic pump, we employ a thermal bubble micropump to enhance the integration of the cytometry chip. The thermal bubble micropump works on the basis that the liquid surrounding the thermal bubble resistor soon evaporates to create a void, which forces the liquid out of the nozzle and propels the liquid flow in the microchannel. The size of thermal bubble resistor is 50 × 40 μm², which is capable of pumping 32 pL per heating, and the flow rate can be changed by adjusting the heating frequency.

TIA can collect and amplify the differential information of cell impedance between the two electrodes. To deploy multiple detection channels, TIA should provide a great anti-noise capacity and occupy a smaller area, particularly for its operational amplifier. Fig. 2(b) depicts the schematic of the two-stage operational amplifier in TIA. The input is made through the differential pair composed of PM1 and PM2, and the operating amplifier reaches enough amplification times. The flicker noise coefficient of NMOS transistors in the CMOS process is significantly higher than the PMOS transistors. To reduce the impact of flicker noise, we select PMOS transistors for the input stage of TIA, and can further reduce noise by increasing their sizes and aspect ratios $\left(\dfrac{W}{L}\right)$. The R−C Miller compensation structure, which is made up of resistor R_C and capacitor C_C connected between the input and output of the second-stage amplifier, compensates for frequency and guarantees that the circuit has enough phase margin to maintain stability while operating. The maximum frequency of the excitation voltage signal is known to be 5 MHz. Therefore, the integrated operational amplifier design must satisfy the unit gain bandwidth requirement of more than 50 MHz. Furthermore, its DC gain needs to be greater than 80 dB in order to increase the precision of the differential amplifier. The open-loop amplitude-frequency and phase-frequency characteristics of the two-stage operational amplifier is displayed in Fig. 2(c), which were acquired by circuit simulation. The DC gain, GBW, and phase margin are 84 dB, 57 MHz, and 68 °, correspondingly, which can meet all the design targets.

As shown in Fig. 2(d), the feedback resistor R_F of TIA can be programmed through a digital interface and is divided into seven levels, ranging from 5−60 kΩ, to accommodate a wide range of input impedances. Meanwhile, the value of R_F can be changed by activating the corresponding switch with the digital control signal. The output ${V}_{1}$ and ${V}_{2}$ of differential TIAs can be calculated by the following formulas:

$${V}_{1}=-\frac{{R}_\rm{F}}{{Z}_{0}+{Z}_{\mathrm{c}\mathrm{e}\mathrm{l}\mathrm{l}}}\cdot \mathrm{A}\mathrm{s}\mathrm{i}\mathrm{n}({\omega }{t}),$$

(4)

$${V}_{2}=-\frac{{R}_\rm{F}}{{Z}_{0}}\cdot \mathrm{A}\mathrm{s}\mathrm{i}\mathrm{n}({\omega }{t}).$$

(5)

To prevent the influence of parasitic parameters from discrete devices and achieve a high degree of integration of cell counting channels, we integrate microchannels for cell counting and TIAs using CMOS-MEMS technology, resulting in a total of 36 integrated channels on the same chip. In addition, the 36 detection channels can operate autonomously, effectively averting a complete failure of the cell counter in the event of a single flow channel blockage, which enhances the overall robustness of the flow cytometry chip. The TIAs and the digital logic circuit require a voltage of 5 V as the supply power, while the thermal bubble micropumps require a voltage of 35 V. Therefore, the CMOS and MEMS modules must be placed far apart to prevent crosstalk between different voltage domains. However, this requires the weak signal on the induction electrodes to travel a long distance in order to reach the TIAs, so we set up the distance between the CMOS and MEMS power cord to 400 μm.

The gating of various detection channels can be accomplished using a logic control circuit. The micropump and TIA are closed when the detection channel is closed correspondingly, which can lower the overall static power consumption of the chip.

2.3 Fabrication, encapsulation and test

As shown in Fig. 3, the fabrication of the 36 channels EIS flow cytometry can be divided into two parts: CMOS front process and MEMS post process.

The CMOS part of the chip including TIA arrays, current bias circuits, and digital logic circuits was fabricated on a 4-metal 2-poly layer using a standard 0.35 μm CMOS process from the XFAB. The MEMS components of the chip were fabricated on 8-inch Pilot Line of SITRI. Its primary processes involve the deposition of tantalum electrodes and heater resistors, followed by the etching of liquid inlets, microchannels, and micropillar arrays in an epoxy resin layer. Ultimately, the nozzle is etched in a dry film layer to create a micropump. As can be seen from Fig. 3, the induction electrodes are directly connected to the top metal layer by through silicon vias (TSV), and is interconnected to the TIA input end through multiple layers of metal, greatly reducing the transmission path of the signal.

The front and back views of EIS flow cytometry chip after packaging are depicted in Figs. 4(a) and 4(b), respectively. Up to 20 μL of cell suspensions can be kept in the customized tank for further cell counts. To stop the liquid pumped from the chip from contacting with the metal wire, 79 PADs of the chip are secured and covered with glue once they have been bonded to the printed circuit board (PCB).

The entire circuits take up approximately 6.7 × 3.6 mm² of space, and from Fig. 4(c), we can see that microfluidics and sensing area of one cell counting channel of 202 × 596 μm². Furthermore, the 216 × 242 μm² differential TIA circuit is composed of two symmetrical switching arrays, feedback resistor arrays, and two-stage operational amplifiers (Fig. 4(d)).

To evaluate the performance of EIS flow cytometry chip, we set up a test platform where electricity and optics are mutually assisted, as shown in Figs. 5(a) and 5(c). First of all, to get exact control over the TIA feedback resistor and the thermal bubble micropump, we utilize a logic analyzer to input the channel gating and circuit configuration data into the shift register array inside the chip. As shown in Fig. 5(b), the excited sinusoidal voltage signal $\mathrm{A}\mathrm{s}\mathrm{i}\mathrm{n}\left({\omega }\mathrm{t}\right)$ for middle electrodes is supplied by the HF2LI lock-in amplifier, which also receives the differential output voltage signal ${V}_{0}$ from the TIAs. The formula for calculating ${V}_{0}$ is as follows:

$${V}_{0}={V}_{1}-{V}_{2}=\frac{{R}_\rm{F}}{{Z}_{0}\left({Z}_{0}+{Z}_{\mathrm{c}\mathrm{e}\mathrm{l}\mathrm{l}}\right)}\cdot \left|{Z}_{\mathrm{c}\mathrm{e}\mathrm{l}\mathrm{l}}\right|\cdot \mathrm{A}\mathrm{s}\mathrm{i}\mathrm{n}\left({\omega }{t}+{\theta }\right).$$

(6)

And V₀ multiplies it with the in-phase signal $\mathrm{A}\mathrm{s}\mathrm{i}\mathrm{n}\left({\omega }{t}\right)$ and the quadrature signal $\mathrm{A}\mathrm{c}\mathrm{o}\mathrm{s}\left({\omega }{t}\right)$ respectively. The high-frequency components in the signal are filtered out by a low-pass filter, and the remaining DC components are ${V}_\rm{re}$ and ${V}_\rm{im}$, as shown below：

$${V}_\rm{re}={V}_{0}\cdot \mathrm{A}\mathrm{s}\mathrm{i}\mathrm{n}\left({\omega }{t}\right)=\frac{1}{2}{C}\left|{Z}_{\mathrm{c}\mathrm{e}\mathrm{l}\mathrm{l}}\right|\cdot \left[-\mathrm{cos}\left(\theta \right)\right],$$

(7)

$${V}_\rm{im}={V}_{0}\cdot \mathrm{A}\mathrm{c}\mathrm{o}\mathrm{s}\left({\omega }{t}\right)=\frac{1}{2}{C}\left|{Z}_{\mathrm{c}\mathrm{e}\mathrm{l}\mathrm{l}}\right|\cdot \left[\mathrm{sin}\left(\theta \right)\right].$$

(8)

Here $C=\rm{A}^{2}\cdot \dfrac{{R}_\rm{F}}{{{Z}_{0}}^{2}}$ can be regarded as a constant. Finally, ${V}_\rm{re}$ and ${V}_\rm{im}$ are sampled by an internal ADC of HF2LI lock-in amplifier and transmitted to the laptop. The impedance information including amplitude $\left|{Z}_{\mathrm{c}\mathrm{e}\mathrm{l}\mathrm{l}}\right|$ and phase $\theta$ of the cell can be extracted. $\left|{Z}_{\mathrm{c}\mathrm{e}\mathrm{l}\mathrm{l}}\right|$ and ${\theta }$ are calculated as follows:

$$\left|{Z}_{\mathrm{c}\mathrm{e}\mathrm{l}\mathrm{l}}\right|=\frac{2}{{C}}\sqrt{{V}_\rm{re}^{2}+{V}_\rm{im}^{2}},$$

(9)

$$\theta =\mathrm{a}\mathrm{r}\mathrm{c}\mathrm{t}\mathrm{a}\mathrm{n}\left|\frac{{V}_\rm{im}}{{V}_\rm{re}}\right|.$$

(10)

Using the oscilloscope, we can view the results of the TIA output voltage. Inverted microscopy, as an auxiliary method in the light field, is used to verify the cell counting accuracy of EIS flow cytometry. DC power is used to provide different stable voltage values.

Polystyrene beads are chosen to assess the counting capabilities of the chip, which are comparable in size and dielectric characteristics to cells. Since blood cells are more difficult to obtain, we chose polystyrene beads with diameters of 7 and 14 μm to substitute to red and white blood cells, to evaluate the counting ability of the chip. To demonstrate cell-counting capability of the designed chip, we additionally test the technical accuracy using MSTO-211H cells whose size is 17 μm.

3. Results and discussions

3.1 Electrical characterization of on-chip TIA

We chose one of the detecting channels for testing to assess the current amplification ability as well as the switching ability of various magnifications of TIA. Using a pipette, 20 μL of phosphate buffered saline (PBS) solution was first poured into the chip liquid storage tank. Next, the intermediate electrode was subjected to a 500 kHz sinusoidal voltage signal with a 0.5 V amplitude. Finally, the logic analyzer was used to switch different feedback resistance gears, and an oscilloscope was used to view the output voltage of TIA. Fig. 6(a) displays the output voltage waveform of TIA with various feedback resistances R_F (5, 10, 20, 30, 40, 50, and 60 kΩ). The mismatch of feedback resistances during manufacturing will impair the magnification accuracy of TIA. During the layout design, we adopted the method of using dummy resistors and cross-matching to reduce the impact of mismatch. The output voltage amplitude of the TIA with various feedback resistors appears in Fig. 6(b), along with the result of its linear fitting. The R² value for the linear regression is 0.99728, indicating a strong linear relationship between the resistance values of different gain gears and the amplitude of the output voltage.

3.2 Classification of polystyrene beads with different particle sizes

As shown in Fig. 7(a), we measured the peak voltage amplitude (I channel) caused by 300 7 μm (red) and 14 μm (blue) polystyrene beads, which were diluted in PBS solution, passing through the induction electrodes, respectively. The excitation voltage signal is set to 500 kHz, the amplitude is 0.5 V, the TIA feedback resistance is 5 kΩ, and the thermal bubble micropump pumps the solution at a speed of 1080 μm/s. The HF2LI impedance analyzer processes the differential analog output voltage received from the TIA by employing modulation, filtering, and sampling. Subsequently, it transmits the processed data to the laptop for further analysis. Fig. 7(b) shows the voltage amplitude histogram of 7 and 14 μm beads measurements, which are recorded in Fig. 7(a). 7 and 14 μm beads result in an average voltage change of 0.0529 and 0.4510 mV, correspondingly. It can also be seen in Fig. 7(b), that for a given size of beads, the amplitude of output voltage has a certain deviation. The standard deviation of response amplitude caused by 7 and 14 μm beads is 0.0132 and 0.0975 mV, respectively. A possible explanation for this can be that the vertical position of a bead in the channel generates a variation in its impedance as it passes through the detecting electrodes. However, there is no overlap between the two sets of data which indicates that the chip can clearly classify different particle sizes (7 and 14 μm) of beads (cells) by the change of voltage magnitude.

3.3 Counting accuracy

An inverted microscope linked to a camera was used to synchronously record beads or cells passing through the detection electrodes to verify electrical measurements. In order to identify beads or cells in the light field passing through the detection electrodes, lower the liquid flow speed to 108 μm/s by adjusting the heating frequency of the thermal bubble micropump heater. The typical transient waveform of a 7 μm bead, which modifies the output voltage (I channel) and causes a peak change of 0.04 mV as it passes through the detection electrodes, as depicted in Fig. 8(a). In the meantime, the microscope images of times A, B, and C, which correspond to the beads entering the top electrode, middle electrode, and bottom electrode, respectively, are also shown in Fig. 8(a).

To evaluate the accuracy of cell counting by EIS flow cytometry chip, the counting outcomes of 300 7 μm (red) and 14 μm (blue) polystyrene beads and 211H cells (green) were compared in the electric field and the light field, respectively. Preparing the beads and cells as a monodisperse bead solution beforehand can lessen the impact of cluster cells on the count. The results of the counting and linear fitting of 7 μm (red) and 14 μm (blue) polystyrene beads and 211H cells (green), respectively, are given in Figs. 8(b)−8(d). The counting results under the optical field are represented by the horizontal coordinate, while the counting results under the electric field are represented by the vertical coordinate. The statistics of the counting accuracy of the EIS flow cytometry chip for three different types of particles are displayed in Table 1. The lowest linear fitting coefficient R² was 0.99994%, and the detection accuracy was higher than 97.6%. These results collectively show that this chip has a high level of precision in beads/cells counting. The presence of a few polystyrene beads or cells joined together in the prepared dispersive single-bead solution was the primary cause of the inaccuracy of the detection results.

4. Conclusion

A synthetic comparison of the performances here demonstrated with respect to what offered by impedance flow cytometers recently reported in the literature^{[19, 20, 23, 27−31]} is shown in Table 2. In this work, a thermal bubble micropump-driven EIS flow cytometry chip with on-chip TIA has been designed and fabricated. Multiple detection channels can operate simultaneously thanks to the combination of thermal bubble micropump, differential coplanar electrodes, and microchannel. Compared to our past research^[20], to mitigate the parasitic effect on individual device wires, the TIAs and the MEMS sensors are integrated on the same chip. Simultaneously, the dimension of the TIA circuit is engineered to align with the microfluidic component, enabling the high-throughput integration of the sensors and detection circuits. The EIS flow cytometry chip was fabricated utilizing CMOS-MEMS technology, enabling the integration of 36 counting channels. In the future, there is potential for expansion to incorporate additional channels, thereby enhancing the counting speed. In this study, the testing procedure focused on evaluating the amplification capability and the adjustable function of the multi-gear feedback resistance of the TIA embedded on the chip. The 7 and 14 μm polystyrene beads solutions were used to test the classification ability and the counting accuracy of the chip. From the obtained results, we conclude that the designed chip can precisely classify and count 7 and 14 μm beads. Finally, the counting accuracy of the chip was also validated using 211H cells, and the counting accuracy reached 98%, marking a substantial improvement compared to prior studies, which affirming the biomedical applications potential of the chip. In addition to TIA, in the future, it is possible to use the CMOS-MEMS process to integrate EIS flow cytometry with lock-in amplifiers and analog-to-digital converters (ADCs) on the same chip, which holds promise for future applications in multi-channel, low-power, high-throughput, and high precision cell counting. The cell counter chip integrating sensing and reading out circuits is poised to address the requirements of POCT.

MSTO-211H is a human lung cancer cell line, MCF-7 is human breast cancer cell, HCM is hypertrophic cardiomyocyte, RBC is red blood cell, WBC is white blood cell, CTC is circulating tumour cell, THP-1 is one of leukocytes cell lines, and MDA-MB-231 and MDA-MB-468 are both breast cell lines.

Acknowledgments

This work was supported by the Key Project of the National Natural Science Foundation of China (Grant No. 82130069).