IMPEDANCE SPECTROSCOPY OF

SINGLE CELLS USING EMBEDDED

AGPDMS ELECTRODES

by

Yuan Wei

A thesis submitted in conformity with the requirements

for the degree of Master of Applied Science

Graduate Department of Mechanical and Industrial Engineering

University of Toronto

© Copyright 2015 by Yuan Wei

ii

Abstract

Impedance Spectroscopy of Single Cells Using Embedded AgPDMS Electrodes

Yuan Wei

Master of Applied Science

Graduate Department of Mechanical and Industrial Engineering

University of Toronto

2015

This thesis presents a microfluidic device with wide channels and embedded AgPDMS

electrodes for measuring the electrical properties of single cells. The work demonstrates the

feasibility of using a large channel design and embedded electrodes for impedance

spectroscopy to circumvent issues such as channel clogging and limited device re-usability.

AgPDMS electrodes were formed on channel sidewalls for impedance detection and cell

characterization. Equivalent circuit models were used to interpret multi-frequency impedance

data to quantify the cytoplasm conductivity and specific membrane capacitance of each cell.

T24 cells were tested to validate the microfluidic system and modeling results. Comparisons

were then made by measuring two leukemia cell lines (AML-2 and HL-60) which were

found to have different cytoplasm conductivity values (0.29 ± 0.15 S/m vs. 0.47 ± 0.20 S/m)

and specific membrane capacitance values (41 ± 25 mF/m2 vs. 55 ± 26 mF/m2).

iii

Acknowledgements

I would like to express my deep gratitude for my advisor, Professor Yu Sun, for his limitless

patience and support during my studies at the University of Toronto. His passion for and

dedication to research is infectious and inspiring, and I can only hope to capture his

enthusiasm with whichever career path I pursue.

I would like to thank all members of the Advanced Micro and Nanosystems

Laboratory (AMNL) for their support and expertise. They humbled me with their knowledge

and insight, motivated me to work hard, and made the lab a wonderful environment I looked

forward to being in every day. I am happy to call these people not just my colleagues, but

also my friends. I would specifically like to acknowledge the members of the microfluidics

subgroup: Mark A. Cachia, John Nguyen, Zhensong Xu, and Yi Zheng. They provided

invaluable contributions every step of the way.

I would like to acknowledge the financial support I have received from the Queen

Elizabeth II Graduate Scholarship in Science and Technology (QEII-GSST) and the NSERC

Canada Graduate Scholarship Master’s Program (CGS M).

I would like to thank my parents for making me the young man I am today. They

taught me the value of discipline and hard work, encouraged me to follow my passion, and

supported me in any and every way they could.

Finally, I would like to thank my lovely partner, Cindy Ge, for absolutely everything

she has done for me. Through every trial and tribulation along this winding journey, she has

been by my side keeping me on the right path. She has been my anchor when I needed to stay

grounded, she has been my wings when I needed to soar. Thank you, Cindy, for being my

cheerleader, my best friend, and my love.

iv

Contents

List of Figures ................................................................................................................................ vi

List of Tables ............................................................................................................................... viii

1. Introduction ............................................................................................................................. 1

1.1 Background ...................................................................................................................... 1

1.2 Electrical characterization techniques .............................................................................. 2

1.3 Electrode design and material ......................................................................................... 6

1.4 Motivation ...................................................................................................................... 12

1.5 Thesis outline ................................................................................................................. 13

2. System Overview ................................................................................................................... 14

3. Fabrication and Experimental Procedure .............................................................................. 17

3.1 Fabrication ...................................................................................................................... 17

3.2 Cell culture ..................................................................................................................... 21

3.3 Equivalent circuit models ............................................................................................... 22

3.4 Signal Analysis ............................................................................................................... 26

4. Results and Analysis .............................................................................................................. 29

4.1 AgPDMS characterization.............................................................................................. 29

4.2 Device characterization .................................................................................................. 33

4.3 Cell characterization ....................................................................................................... 36

5. Conclusion ............................................................................................................................. 43

6.1 Future directions ............................................................................................................. 43

Bibliography ................................................................................................................................. 45

Appendix ....................................................................................................................................... 50

v

A Previous electrode designs ................................................................................................. 50

B AgPDMS photopatterning procedure ................................................................................ 51

vi

List of Figures

Figure 1.1: Different mechanisms for trapping cells for impedance spectroscopy. (a) ROT uses a

rotating electric field. (b) A hydrodynamic trap. (c) Multiple traps to increase throughput. ..........4

Figure 1.2: Various metal and metal-chloride electrode designs for impedance spectroscopy. (a)

Parallel Au electrodes. (b) Planar Pt electrodes. (c) Parallel Pt electrodes. (d) Plug-in Ag/AgCl

electrodes. ........................................................................................................................................6

Figure 1.3: (a) Planar Ag/AgCl electrodes. (b) PGEs for size-based impedance measurements of

CTCs. ...............................................................................................................................................9

Figure 1.4: Conductive PDMS composites. (a) Ag-coated Cu flakes mixed with PDMS reached

conductivities > 104 S/m. (b) CPDMS composite has been used in microfluidic valve, pump, and

mixer applications. (c) AgPDMS-based microheater. (d) AgPDMS-based micropressure sensor.

(e) AgPDMS electrodes used for droplet detection. ......................................................................11

Figure 2.1: Schematic diagram of the single cell measurement microfluidic system. AgPDMS

electrodes are embedded on either side of the microchannel. Impedance at 8 frequencies is

measured as cells flow through the microchannel past the electrodes...........................................16

Figure 3.1: (a) Process flow of device fabrication using soft lithography. (b) Mask/pattern for

SU-8 master. The areas exposed to the UV light become insoluble and form features. (c)

Mask/pattern for AZ P4620 mold. The areas exposed to the UV light become soluble and leave a

mold. ..............................................................................................................................................20

Figure 3.2: (a) SU-8 master and AZ P4620 mold prior to applying AgPDMS. (b) AgPDMS

electrodes in a parallel configuration embedded into the PDMS channel sidewalls. ....................21

Figure 3.3: (a) The device without a cell in the microchannel, the double layer effect is modeled

as an ideal capacitor. (b) Still without a cell, the double layer effect is modeled as a CPE. (c) The

device with a cell, which is modeled as two capacitors, representing the membrane capacitance,

in series with a resistor, representing the cytoplasm resistance. ....................................................25

Figure 3.4: (a) Experimental data showing peak events in a 2 second timeframe. Absolute

impedance values have been translated (and are not accurate) for the illustrative purpose of

showing data from all 8 frequencies. (b) Peak finding algorithm in MATLAB. The red line

vii

represents the baseline value, the green line represents the cut-off threshold, and the red circles

highlight peaks that have been identified.......................................................................................27

Figure 4.1: Electric field of 69 wt % Ag concentration simulated in COMSOL...........................30

Figure 4.2: SEM images showing the AgPDMS electrodes. (a) Device bottom and sidewall with

69 wt % Ag. (b) Device bottom and sidewall with 85 wt % Ag. (c) Surface roughness of bottom

and sidewall with 85 wt % Ag. (d) Top-down overview of 85 wt % Ag device. ..........................32

Figure 4.3: Conductivity of AgPDMS with varying Ag wt %. .....................................................33

Figure 4.4: Impedance profile of channels without cells along with curve fitting. (a) The fit with

the EDL represented by an ideal capacitor. (b) The fit with the EDL represented by a CPE. ......34

Figure 4.5: Device parameters over a 3 day experiment period. ...................................................35

Figure 4.6: Impedance values for 10 µm beads at 11 kHz ............................................................37

Figure 4.7: 8 frequency impedance data with 5 particles detected in a 1 second timeframe. Based

on 11 kHz peak height data, cell events were identified from non-cell events. ............................39

Figure 4.8: Scatter plot of T24 electrical properties (n = 740). Colour represents density

distribution. ....................................................................................................................................40

Figure 4.9: AML-2 and HL-60 electrical properties. HL-60 was found to have higher cytoplasm

conductivity and specific membrane capacitance. .........................................................................41

viii

List of Tables

Table 3.1: Fabrication of alignment markers ................................................................................ 18

Table 3.2: Fabrication of SU-8 seeding and 50 µm feature layers ............................................... 18

Table 3.3: Fabrication of AZ P4620 layer .................................................................................... 19

Table 3.4: AML-2 and HL-60 electrical properties ...................................................................... 22

1. Introduction

1

1. Introduction

1.1 Background

The cell is the basic functional unit of all living organisms, and consists of a cytoplasm

enclosed by a membrane. The cell membrane is a selectively permeable structure that

physically contains intracellular components and separates them from the extracellular

environment. It consists of a phospholipid bilayer embedded with specific proteins that

facilitate the passive and active transport of materials into and out of the cell. The cytoplasm

is composed of the cytosol (the intracellular fluid) and the organelles (such as mitochondria

and the nucleus). The cytosol is made up of various salts and organic molecules dissolved in

water, and comprises the majority of the cytoplasm. Each cell type possesses distinct

biochemical and biophysical properties that provide it its specific physiological functions and

allow it to adapt to its environment.1,2 As such, cells have the potential to be uniquely

identified based on their biochemical and biophysical characteristics.

The identification of cell types based on their intrinsic properties is relevant for both

academic and practical purposes. One of the most common and useful applications of cell

characterization is a complete blood count, which is a valuable clinical tool that can lead to

the diagnosis of certain conditions, such as leukemia and anemia.3 The current gold standard

in cell characterization involves the use of flow cytometry, in which cells are stained with

fluorescent-labeled antibodies and flowed through a detection system.4 The flow is controlled

in a manner that forces the cells to line up in a single file. The single cells are exposed to a

laser, and depending on the fluorescent-labeled antibodies selected, will emit light at a

specific wavelength, that can be used to identify the cells. Flow cytometers are highly

versatile and accurate, but they are also generally limited in accessibility (hospital/clinical

settings only), highly operator skill-dependent in both performing measurements and

analyzing data, and expensive. Therein lies potential in a scaled-down, label-free technique

for cell characterization and identification where cost, resources, and user expertise may be

limited. Microfluidics has emerged as a field that boasts advantages such as biocompatibility,

1. Introduction

2

low sample volume, simple sample preparation, and high throughput in cell-related

experiments. As such, the development of microfluidic technologies for single cell

characterization has been of interest.

1.2 Electrical characterization techniques

There have been many investigations with regards to the characterization of cells based on

their electrical properties, specifically cytoplasm conductivity and specific membrane

capacitance (the capacitance per unit area of the cell membrane).5 Höber pioneered the work

in the 1910s when he reported the conductivity of red blood cells (RBCs).6–8 In the mid-

1920s, Fricke used a mathematical model based on Maxwell’s mixture theory to calculate the

specific membrane capacitance of canine RBCs.9–11 Schwan proposed the single shell model

in the 1950s, where the cell is modeled as a single shelled sphere: the cell membrane is

represented as a thin dielectric shell, which contains the homogeneous resistive cytoplasm.12

Different cells will have different protein expressions and lipid bilayer compositions, as well

as different ion concentrations and nucleus-to-cytoplasm volume ratios, which will affect the

specific membrane capacitance and cytoplasm conductivity, respectively.13,14 Thus,

characterizing these two electrical parameters can lead to identification of the cell.

Different electrical characterization techniques were developed for single cell

experiments. The first high-speed electrical measurement technique on single cells was the

Coulter counter in the 1950s, where it was demonstrated that a direct current (DC)

impedance signal could be used to count and provide size information of each cell.15 In the

1970s, Hoffman and Britt developed a system that used a high frequency signal simultaneous

to a low frequency signal.16 This enabled it to detect both resistive and capacitive changes

caused by the cells, which reflected cytoplasm conductivity and specific membrane

capacitance. The patch clamp technique was developed in 1976 and was used to measure

specific membrane capacitance by aspirating a patch of the cell membrane into a

micropipette.17 However, this technique is complex in nature, typically requiring tens of

minutes per cell and high user expertise. In recent years, more efforts have been directed

towards microfluidic technologies for single cell electrical characterization.

1. Introduction

3

Electrorotation (ROT) was a technique developed in 1982 that uses the forces exerted

by a rotating electric field to rotate the cell.18 ROT is similar to dielectrophoresis (DEP) in

that it depends on a non-uniform field, but it is more often used for single cell electrical

characterization since it depends only on a rotating electric field, and not on a varying

electric field amplitude.19 In contrast, DEP requires a constant electric field gradient, which is

more challenging to implement experimentally than a constant electric field amplitude. ROT

experiments typically use four electrodes to apply sine waves in phase quadrature, as seen in

Figure 1.1 (a).20,21 Cells are placed in the centre of the electrodes, the signal frequency is

varied from kHz to MHz, and the ROT spectra is curve fit to determine cytoplasm

conductivity, cytoplasm permittivity, and specific membrane capacitance.22 While ROT has

been successfully used to measure intrinsic electrical properties of cells, it is not without its

drawbacks. ROT suffers from limited throughput, requiring approximately 30 minutes per

cell.20 In addition, low conductivity buffer must be used since it is difficult to achieve

efficient rotation in a high conductivity buffer, which may alter the electrical properties of

the cells.

Impedance spectroscopy is a technique where an excitation signal is applied to a cell

at multiple frequencies. Initial designs involved trapping the cell while measuring the

impedance response as a function of frequency. These designs used a variety of trapping

mechanisms, such as hydrodynamics23, pressure24, and DEP.25 The hydrodynamic trapping

mechanism is shown in Figure 1.1 (b), where human cervical epithelioid carcinoma cells are

tested. By curve fitting the impedance spectra to an equivalent circuit model, the electrical

properties can be determined. However, the impedance data is highly dependent on the cell-

trapping mechanism, the size of the electrodes, and the size of the cell. Like patch clamping

and ROT, this technique has a low throughput due to the slow trapping (and subsequent

releasing) procedure. The throughput can be increased by fabricating an array of traps, like

the one shown in Figure 1.1 (c), but the electrodes must then be cleverly designed to allow

multiplexed impedance measurements from multiple electrodes.26

1. Introduction

4

Figure 1.1: Different mechanisms for trapping cells for impedance spectroscopy. (a) ROT uses a

rotating electric field.21 (b) A hydrodynamic trap.23 (c) Multiple traps to increase throughput.26

1. Introduction

5

High-speed microfluidic designs, capable of measuring tens to hundreds of cells per

second, have been developed that use impedance spectroscopy to electrically probe single

cells as they flow through a small microchannel (in the range of the cell size) where the

electric field lines are concentrated. The pioneering device in 1999 used parallel integrated

gold (Au) electrodes situated on either side of the microchannel and could detect impedance

differences in various blood cells, seen in Figure 1.2 (a).27 Co-planar28 and parallel29

platinum (Pt) electrode designs have also been used to theoretically and experimentally

determine cytoplasm conductivity and specific membrane capacitance, shown in Figure 1.2

(b) and (c). Two sets of embedded Pt electrodes, one for sensing and one for reference, have

been tested in providing a more accurate differential measurement of the impedance of the

cell compared to the baseline, and have been used to characterize leukocyte sub-

populations.30,31,32 The embedded electrode design has been used to probe single cells at high

frequency (> 1 MHz), which is necessary to measure intracellular characteristics.12 However,

pure metal electrodes exhibit a high electrical double layer (EDL) effect, caused by the

interaction at the interface of the metal electrodes and the ionic solution, which can greatly

affect low frequency impedance data.27

Several devices have used plug-in Ag/AgCl electrodes have been implemented to

minimize the EDL, allowing for lower frequency measurements and more accurate cell

sizing, though the signal is limited to a sub-MHz frequency range due to stray capacitance

effects.33–35 These devices, demonstrated in in Figure 1.2 (d), implement a channel that is

smaller than the cell to achieve the necessary sensitivity since they will have a much higher

basal impedance compared to embedded electrodes. They have been used to characterize the

electrical properties of cancer cell lines. Though each high-speed microfluidic design has

various advantages and disadvantages, they all dictate that the detection microchannel be

close to the size of the cell for sensitivity requirements. These small (< 20 µm) channels are

subject to practical issues, such as clogging of the device due to debris. Thus, it is of interest

to investigate the feasibility of large channel designs in measuring single cell electrical

properties.

1. Introduction

6

Figure 1.2: Various metal and metal-chloride electrode designs for impedance spectroscopy. (a)

Parallel Au electrodes.27 (b) Planar Pt electrodes.28 (c) Parallel Pt electrodes.29 (d) Plug-in Ag/AgCl

electrodes.33

1.3 Electrode design and material

Various electrode materials and designs have been investigated for single cell impedance

spectroscopy applications. Plug-in electrodes compared to embedded electrodes, while

simpler to fabricate, show a much higher basal impedance (MΩ vs. tens of kΩ) due to being

much farther apart, which means the dimensions of the microchannel must be smaller to

increase the signal to noise ratio.30,31,33–35 In addition, at higher frequencies, the plug-in

electrode design is more susceptible to stray capacitance effects when the electric field lines

undesirably obviate from the microchannel and the cell inside. This means the high

frequency range required to probe the internal properties of the cell is more limited, affecting

the accuracy of measurements of electrical properties. Thus, embedded electrodes are

preferable compared to plug-in electrodes.

1. Introduction

7

The next decision with regards to electrode design is determining whether planar

electrodes or parallel electrodes are more suitable for impedance spectroscopy. Gawad et al.

demonstrated a microfluidic device with planar electrodes capable of differentiating different

sized latex beads, as well as RBCs from ghost RBCs (empty membranes).28 However, the

planar electrode configuration generated a non-uniform electric field, which had a

considerable effect on the variation in the impedance amplitude due to particles flowing at

different positions in the channel. This configuration requires an accurate mapping of the

electric field distribution to model the impedance spectrum of a single cell. In contrast, the

impedance signal was found to be less dependent on particle position in a parallel electrode

configuration since the electric field was more uniform.36 Therefore, the embedded parallel

configuration constitutes the most suitable electrode design.

The selection for the material of the electrodes is important as electrically stable,

biocompatible electrodes are desirable. That is, the electrodes must be non-degradable when

in contact with the saline buffer solutions typically used to bathe cells during experiments,

and must not contaminate or compromise the integrity of the cells for the purpose of

accuracy in not altering the cell properties and in case they are to be collected for further

experimentation. Solid metal electrodes were the first material to be used in single cell

impedance spectroscopy applications due to their well-defined electrical properties.27

Fabrication of such electrodes requires electroplating or evaporation as well as alignment of

separate pieces for bonding.27,37 These electrodes were capable of performing measurements

at frequencies as high as 10 MHz, but could only take measurements as low as 500 kHz,

compromising size data. This is due to solid metal electrodes exhibiting a high EDL, which

can have a substantial effect on low frequency impedance data.27

Ag/AgCl is non-polarizable and highly conductive, making it an ideal choice for the

electrode material and the most common type of electrode used in research and industry.38

However, the miniaturization and fabrication of integrated Ag/AgCl electrodes is non-trivial

since most Ag/AgCl thin films have poor stability due to dissolution of the AgCl, leading to

mixed potentials at the electrode/solution interface.39 Embedded Ag/AgCl electrodes have

been fabricated, as seen in Figure 1.3 (a), but are planar, while the parallel electrode

configuration is preferred for generating a uniform electric field.39,40 In addition, these

1. Introduction

8

electrodes have a complex fabrication process as they require alignment and bonding of

separate pieces, as opposed to a single solid structure.40

1. Introduction

9

Figure 1.3: (a) Planar Ag/AgCl electrodes.39(b) PGEs for size-based impedance measurements of

CTCs.41

There have been efforts to fabricate conducting composite electrodes to minimize the

EDL. For example, polyelectrolyte gel electrodes (PGEs), shown in Figure 1.3 (b), have been

shown to detect circulating tumour cells (CTCs) using size-based direct current (DC)

1. Introduction

10

impedance measurements.41 However, they are limited to differentiating particles with

significant size differences, as they have not been tested with alternating current (AC) signals.

Many microfluidic devices are fabricated using polydimethylsiloxane (PDMS) due to its

favourable physical and chemical properties, such as flexibility, nontoxicity, biocompatibility,

chemical inertness, gas permeability, and transparency.42 Fabricating PDMS-based devices is

also simpler and less costly than silicon or glass based devices, which allows for rapid design

iterations and prototyping. Naturally, there has been interest in developing PDMS-based

conducting composite electrodes, as integration of these electrodes into PDMS-based

microfluidic chips would be straightforward. These would have the mechanical properties of

PDMS while still being appropriately conductive for impedance measurement. It should be

noted that PDMS and PDMS composites can be fabricated via screen printing or even

photopatterning, compared to more intensive methods for pure metal electrodes, such as

sputtering or evaporation.37,43–46

PDMS-metal composites are viable solutions due to percolation theory, which states

that there is a threshold of conductive particle concentration at which a continuous chain will

be formed, resulting in macroscopic conductivity.47 This threshold depends on a number of

factors, such as the size of the particles, the geometry of the particles, and the mixing

procedure. Increasing the conductive particle concentration beyond this threshold will lead to

a significant decrease in resistivity. Thus, given sufficient concentration of the conductive

filler, the composite can be rendered highly conductive. Silver (Ag) coated copper (Cu)

flakes were mixed with PDMS to achieve conductivities greater than 104 S/m at

concentrations between 75 wt % and 85 wt %.48 Carbon PDMS (CPDMS) composites, seen

in Figure 1.4 (a), were made from carbon black powder, and were found to be conductive

above 10 wt %, although the conductivity was much lower (on the order of 10 S/m).44,49

CPDMS has been tested in several applications, including microfluidic valves50, pumps51,

and mixers.52

Silver has been of particular interest in such composite electrodes. Compared to other

metals, such as titanium (Ti) and copper, Ag yields the lowest interfacial impedance, which

is relevant in minimizing signal distortion and performing with an adequate signal-to-noise

ratio (SNR).53,54 Silver PDMS (AgPDMS) is also much more conductive than CPDMS,

1. Introduction

11

reaching conductivities greater than 104 S/m.44 AgPDMS electrodes have previously been

used in a variety of applications. AgPDMS have been used in temperature-dependent devices,

such as a microheater, seen in Figure 1.4 (b), and a thermochromics display.45,55,56 AgPDMS

conductivity also varies as a function of strain, showing potential as a micropressure sensor,

seen in Figure 1.4 (c).44,46 Finally, AgPDMS has relevantly been used in droplet detection

applications, where impedance signals vary based on droplet size, seen in Figure 1.4 (d).57

Despite these numerous applications, AgPDMS electrodes have not yet been demonstrated to

perform single cell electrical characterization.

Figure 1.4: Conductive PDMS composites. (a) Ag-coated Cu flakes mixed with PDMS reached

conductivities > 104 S/m.48 (b) CPDMS composite has been used in microfluidic valve, pump, and

mixer applications.49 (c) AgPDMS-based microheater.55 (d) AgPDMS-based micropressure sensor.46

(e) AgPDMS electrodes used for droplet detection.57

1. Introduction

12

1.4 Motivation

There are several microfluidic devices designed to measure the intrinsic electrical properties

of single cells, specifically cytoplasm conductivity and specific membrane capacitance.

Characterizing these properties can lead to identification of the cell, which is useful in

applications such as the complete blood count. Microfluidic devices typically perform this

characterization by using impedance spectroscopy to probe cells as they pass through a small

channel. Some of these devices use embedded metal electrodes, which are able to probe cells

at a high frequency (> 1 MHz), but exhibit a high EDL which can distort low frequency

impedance data.27,31 Other devices use plug-in Ag/AgCl electrodes to avoid the EDL and

obtain more accurate cell sizing, but are limited to sub-MHz frequency range due to stray

capacitance effects.33 Though each device has various advantages and disadvantages, they all

dictate that the detection microchannel be close to the size of the cell for sensitivity

requirements.

In this thesis, I propose a wide channel, AgPDMS embedded electrode microfluidic

chip capable of measuring cytoplasm conductivity and specific membrane capacitance of

single cells. A large 100 µm channel design is implemented to overcome practical issues

such as clogging of the channel and being restricted to one-time-use only. The AgPDMS

composite is selected as the material for integrated electrodes for impedance detection that

have adequate sensitivity for the electrical characterization of cells. Cell lines are to be tested

to both verify the feasibility of the wide channel design for data collection, and to validate

the characterization and differentiation of the cells based on electrical properties. This will be

accomplished by fitting multi-frequency impedance data to equivalent circuit models to

characterize both the device and single cell electrical properties. Specific objectives of the

thesis include:

Designing, fabricating, and characterizing a microfluidic device for single cell

electrical characterization

Evaluating if a large channel design is viable for cell characterization while avoiding

clogging

Determining and characterizing the proper Ag concentration for AgPDMS electrodes

1. Introduction

13

Evaluating the appropriate equivalent circuit models for representing the device with

and without a cell

Characterizing cell lines based on their cytoplasm conductivity and specific

membrane capacitance

1.5 Thesis outline

An overview of the subsequent chapters is as follows: Chapter 2 provides a system overview

of the microfluidic chip that measures single cell electrical properties. Chapter 3 details the

fabrication procedures and experimental methods. Chapter 4 presents the experimental

results and discussion on the characterization of the device and the measurement of single

cell electrical properties. Finally, Chapter 5 concludes the thesis with a summary of

contributions as well as potential directions for future work.

2. System Overview

14

2. System Overview

Figure 2.1 shows a schematic diagram of the single cell measurement system. The

microfluidic chip is made of PDMS, with parallel AgPDMS electrodes embedded on either

side of a 100 µm microchannel, bonded to a standard 1” x 3” glass slide. All device features,

including the microchannel and the AgPDMS electrodes, have a height of 50 µm, and the

AgPDMS electrodes are 100 µm in width. Thus, the measurement region is 100 x 100 x 50

µm3. Cells are typically around 10 µm in diameter, which means the volume fraction of the

cell (the ratio of the volume to the cell to the volume of the electrode measurement region) is

< 1%. Previously reported devices had a maximum detection region of 20 x 20 x 20 µm3, and

thus a cell occupied a volume fraction of > 6%.

The microfluidic chip interfaces with the HF2IS Impedance Spectroscope (Zurich

Instruments). The impedance spectroscope employs a function generator and 8 demodulators

to generate a sinusoidal excitation input voltage at 8 different frequencies simultaneously.

These frequencies (11 kHz, 101 kHz, 401 kHz, 701 kHz, 1.01 MHz, 2.01 MHz, 4.01 MHz,

6.01 MHz) were chosen based on the SNR. The input voltage was given an amplitude of 0.4

Vp-p to optimize the SNR without risk of electroporation of the cells.58 Current in the

microchannel is pre-amplified and demodulated at a sampling rate of 14.4 kHz for each

frequency component.

The device is loaded with phophaste-buffered saline (PBS) with 1% w/v bovine

serum albumin (BSA) to maintain cell integrity and to reduce cell adhesion on the channel

walls during experiments. Cells are loaded in the inlet using a needle and are pulled to the

outlet at a pressure of 100 Pa using a syringe. The pressure was selected to minimize

coincident events (simultaneous passage of multiple cells through the channel). Flow

focusing branches generate sheath flow and are used to ensure cells travel down the center of

the microchannel as they pass by the measurement region to minimize impedance distortion.

Sheath flow has also been known to prevent cell adhesion to the device, minimizing sample

loss.41 Additional PBS and cells can be loaded in the inlet without stopping the experiment.

2. System Overview

15

The input voltage signals from the impedance spectroscope generate an electric field

in the microchannel. As cells pass through the microchannel, they perturb the electric field

and affect the current (in turn, affecting the impedance). The impedance difference at each

frequency depends on the size, shape, and electrical properties of the cell. Thus, the

impedance of cells at each of the 8 frequencies is measured, creating an impedance profile.

This impedance profile is fitted to equivalent circuit models to determine cytoplasm

conductivity and specific membrane capacitance for the cell.

All experiments were conducted with the same device, as the large width of the

channel prevented any clogging from debris. Experiments were conducted over 3 days (one

day per cell line), with de-ionized (DI) water being flushed through the device at the end of

each experiment for cleaning purposes.

2. System Overview

16

Figure 2.1: Schematic diagram of the single cell measurement microfluidic system. AgPDMS

electrodes are embedded on either side of the microchannel. Impedance at 8 frequencies is measured

as cells flow through the microchannel past the electrodes.

3. Fabrication and Experimental Procedure

17

3. Fabrication and Experimental Procedure

3.1 Fabrication

The microfluidic chip was fabricated using standard lithographic techniques. The masks for

the device were designed in AutoCAD (Autodesk, Inc., USA) and were printed as

transparencies. Figure 3.1 (a) shows the process flow of the fabrication procedure. Alignment

of features on the photoresist layers was accomplished by first patterning a chromium layer

on a glass slide. Details on alignment marker fabrication are shown in Table 3.1. The SU-8

negative photoresist (MicroChem, Newton, MA, USA) was used both as a seeding layer and

to create a 50 µm high master. Details on fabricating the SU-8 layers are shown in Table 3.2,

and the mask for the SU-8 is shown in Figure 3.1 (b). The AZ P4620 positive photoresist

(Capitol Scientific, Austin, TX, USA) was then spin-coated and patterned to create a 50 um

high mold for the AgPDMS electrodes. Details on fabricating the AZ P4620 layer are shown

in Table 3.3, and the mask for the SU-8 is shown in Figure 3.1 (c). The master with the SU-8

and AZ P4620 layers is shown in Figure 3.2 (a).

Ag particles (Sigma-Aldrich, St. Louis, MO, USA) were mixed with PDMS

(Ellsworth Adhesives, Stoney Creek, ON, Canada) at 85 wt % using a mortar and pestle. The

AgPDMS composite was then spread into the mold created by the AZ P4620 positive

photoresist, with excess AgPDMS being removed using a razor blade to leave the desired

pattern. The master was baked at 70 °C for 30 minutes to cure the AgPDMS, and then

washed in acetone to dissolve the AZ P4620 layer, followed by methanol and DI water.

PDMS was made with the standard 10:1 mixing ratio of base to curing agent. PDMS was

then poured onto the mold and cured on the hotplate at 150 °C for 10 minutes. The entire

PDMS and AgPDMS structure was peeled off the SU-8 master. Due to the adhesive

properties of PDMS, AgPDMS was found to bond without issue to PDMS, creating a single

unit. The structure was washed in acetone, methanol, and DI water, dehydrated on a hotplate

at 150 °C for 10 min, and then O2 plasma bonded to a clean glass slide. The final

microchannel was 100 µm wide and 50 µm high, and the AgPDMS electrodes have a length

3. Fabrication and Experimental Procedure

18

of 100 µm along the sidewall and are also 50 µm high. The microchannel and electrodes are

partially pictured in Figure 3.2 (b).

Table 3.1: Fabrication of alignment markers

Step Procedure

1 Starting with pre-coated chromium slides, clean with acetone, methanol, and DI water

2 Dehydrate glass slides on hotplate at 150 °C for 30 min

3 Use P-20 (20% HMDS) primer to liquid prime entire slide

4 Spin: step 1 – 500 rpm, 10 s, 1 acl; step 2 – 3000 rpm, 30 s, 8 acl

5 Pour photoresist S1811 in middle of slide

6 Spin: step 1 – 500 rpm, 10 s, 1 acl; step 2 – 3000 rpm, 30 s, 8 acl

7 Bake slide on hotplate at 100 °C for 2 min to remove solvent

8 Place slide and mask in mask aligner. Soft contact UV exposure for 6 s

9 Develop slide in MF-321 for 2 min. Rinse in DI water and dry with N2 gun

10 Hard bake slide on hotplate at 100 °C for 1 min

11 Etch chromium layer in CR-2 for 1 to 2 min. Rinse in DI water and dry with N2 gun

12 Develop in AZ-300T for about 5 min. Rinse in DI water and dry with N2 gun

Table 3.2: Fabrication of SU-8 seeding and 50 µm feature layers

Step Procedure

1 Dehydrate glass slides on hotplate at 150 °C for 30 min

2 Pour SU-8-5 on entire slide

3 Spin: step 1 – 500 rpm, 10 s; 1 acl, step 2 – 3000 rpm, 30 s, 3 acl

4 Pre-exposure bake: 65 °C for 1 minute, 95 °C for 3 min

5 Place slide in mask aligner. Flood exposure for 6 s

6 Post-exposure bake: 65 °C for 1 minute, 95 °C for 1 min

7 Hard bake at 175 °C for 2 hours

8 Pour SU-8-25 on entire slide

3. Fabrication and Experimental Procedure

19

9 Spin: step 1 – 500 rpm, 10 s; 1 acl, step 2 – 1000 rpm, 30 s, 3 acl

10 Pre-exposure bake: 65 °C for 5 min, 95 °C for 15 min

11 Place slide and mask in mask aligner. Soft contact exposure for 12 s

12 Post-exposure bake: 65 °C for 1 min, 95 °C for 4 min

13 Develop in SU-8 developer for several min

14 Hard bake at 175 °C for 2 hours

Table 3.3: Fabrication of AZ P4620 layer

Step Procedure

1 Dehydrate glass slides on hotplate at 150 °C for 30 min

2 Pour AZ P4620 on entire slide

3 Spin: step 1 – 300 rpm, 2 s, 1 acl; step 2 – 0 rpm, 10 s; step 3: 650 rpm, 60 s, 3 acl

4 Bake: 110 °C for 5 min

5 Pour AZ P4620 on entire slide

6 Spin: step 1 – 300 rpm, 2 s, 1 acl; step 2 – 0 rpm, 10 s; step 3: 700 rpm, 60 s, 3 acl

7 Bake: 110 °C for 5 min

8 Hydrate slides by placing in dark in petri dish with damp cleanroom wipe for 2 hours

9 Place slide and mask in mask aligner. Soft contact exposure for 210 s

10 Develop in 1:4 AZ 400K developer for several minutes

3. Fabrication and Experimental Procedure

20

Figure 3.1: (a) Process flow of device fabrication using soft lithography. (b) Mask/pattern for SU-8

master. The areas exposed to the UV light become insoluble and form features. (c) Mask/pattern for

AZ P4620 mold. The areas exposed to the UV light become soluble and leave a mold.

3. Fabrication and Experimental Procedure

21

Figure 3.2: (a) SU-8 master and AZ P4620 mold prior to applying AgPDMS. (b) AgPDMS electrodes

in a parallel configuration embedded into the PDMS channel sidewalls.

3.2 Cell culture

T24 (bladder carcinoma), AML-2 (acute myeloid leukemia), and HL-60 (human

promyelocytic leukemia) cell lines were selected for testing due to availability and previous

literature data, summarized in Table 3.4.33 AML-2 and HL-60 are both leukemia cell lines

and have been shown to be of similar size; thus any differences in impedance can be

attributed to differences in electrical properties.33 T24 cells were purchased from the

American Type Culture Collection (ATCC, Manassas, VA, USA), and were cultured in

ATCC-formulated McCoy’s 5a modified medium supplemented with 10% fetal bovine

serum (FBS) (Life Technologies, Grand Island, NY, USA) and 1% penicillin (Life

Technologies, Grand Island, NY, USA). AML-2 and HL-60 were obtained from Mount Sinai

Hospital (Toronto, ON, Canada), and were cultured in Dulbecco’s Modified Eagle’s Medium

(Sigma-Aldrich, St. Louis, MO, USA) supplemented with 10% FBS and 1% penicillin. All

cell lines were incubated at 37 °C with 100% humidity and 5% CO2. Prior to experiments,

3. Fabrication and Experimental Procedure

22

cell lines were centrifuged and re-suspended in PBS (Life Technologies, Grand Island, NY,

USA) with 1% w/v BSA (Life Technologies, Grand Island, NY, USA) at room temperature.

Table 3.4: AML-2 and HL-60 electrical properties33,35

Cell line σcyto (S/m) Cmem (mF/m2)

T24 N/A 47.0 ± 5.1

AML-2 0.62 ± 0.10 12.0 ± 1.44

HL-60 0.76 ± 0.12 14.5 ± 1.75

3.3 Equivalent circuit models

It has been shown that Maxwell’s mixture theory can be used to model dielectric properties

of a single particle positioned between two parallel electrodes.11 As derived elsewhere30, the

equivalent complex permittivity of a mixture containing a particle in a suspending medium is

described as:

𝜀𝑖𝑥 = 𝜀𝑒𝑑1+2𝜙𝐶𝑀

1−𝜙𝐶𝑀 (1)

where 𝜀𝑒𝑑is the complex permittivity of the medium, 𝜙 is the volume fraction (the ratio of

the volume of the cell to the volume of the electrode measurement region), and 𝑓𝐶𝑀 is the

Clausius-Mossotti factor. The volume fraction is given by:

𝜙 =4

3𝜋𝑟3

𝑙𝑤ℎ (2)

where r is the radius of the cell, l = 100 µm is the length of the electrode along the channel

(parallel to the direction of flow), w = 100 µm is the width of the channel, h = 50 µm is the

height of the electrode and channel. The Clausius-Mossotti factor is calculated as:

𝑓𝐶𝑀 =𝑝−𝑚𝑒𝑑

𝑝+2𝑚𝑒𝑑 (3)

where 𝜀𝑒𝑑is the complex permittivity of the particle. From Eq. 1 and based on the single

shell model developed by Schwan12, the complex permittivity of a cell becomes:

3. Fabrication and Experimental Procedure

23

𝜀𝑒𝑙𝑙 = 𝜀𝑒𝑚

𝛾3+2(𝑐𝑦𝑡𝑜−𝑚𝑒𝑚

𝑐𝑦𝑡𝑜+2𝑚𝑒𝑚)

𝛾3−(𝑐𝑦𝑡𝑜−𝑚𝑒𝑚

𝑐𝑦𝑡𝑜+2𝑚𝑒𝑚)

(4)

where 𝛾 =𝑟

𝑟−𝑑, d is the thickness of the cell membrane, 𝜀𝑦𝑡𝑜 is the complex permittivity of

the cytoplasm, and 𝜀𝑒𝑚 is the complex permittivity of the cell membrane. Maxwell’s

mixture theory applies for particles of low volume fractions in a uniform electric field, as is

the case for an ideal parallel plate electrode configuration. However, the geometry of the

system (the distance between electrodes is similar to the dimensions of the electrode) results

in the electric field not being perfectly uniform due to fringe effects. Thus, a geometric

correction factor must be applied when calculating the complex impedance of the system.

This geometric factor was determined to be29:

𝐺𝑓 = ℎ𝜅 = ℎ𝐾(𝑘)

𝐾′(𝑘) (5)

where K is the complete elliptic integral with modulus 𝑘 = tanh (𝜋𝑙

2𝑤) = tanh (

𝜋

2). From this,

the complex impedance of the system is given by30,31:

𝑍𝑚𝑖𝑥 =1

𝑖𝜔𝑚𝑖𝑥𝐺𝑓 (6)

Given the geometrical parameters of the system (i.e. electrode and channel

dimensions), the system may be represented by an equivalent circuit model.30,59 Two

equivalent circuit models were used, one of the system without a cell, and one of the system

with a cell. The microchannel is filled with PBS, which can be represented as a simple

parallel RC circuit with an equivalent resistance RCh in parallel with a capacitance CCh. The

impedance of the channel/PBS medium is given by:

𝑍𝐶ℎ(𝜔) =𝑅𝐶ℎ

1+𝑖𝜔𝑅𝐶ℎ𝐶𝐶ℎ (7)

A cell is represented as a resistor Rcyto representing the cytoplasm in series with two

capacitors Cmem representing the cell membrane. The resistance of the membrane is typically

extremely high compared to the reactance of the membrane and is thus ignored. Similarly,

the reactance of the cytoplasm is negligible compared to the resistance of the cytoplasm. At

low frequencies, the capacitance from the cell membrane dominates the impedance. At high

3. Fabrication and Experimental Procedure

24

frequencies, the capacitance from the membrane approaches zero, allowing measurement of

the resistive cytoplasm. When the cell is in the channel, its impedance is in parallel with that

of the channel. The total impedance of the cell-channel system is given by:

𝑍𝑠𝑦𝑠(𝜔) =𝑅𝐶ℎ(1+𝑖𝜔𝑅𝑐𝑦𝑡𝑜𝐶𝑚𝑒𝑚)

𝑖𝜔𝑅𝐶ℎ𝐶𝑚𝑒𝑚+(1+𝑖𝜔𝑅𝑐𝑦𝑡𝑜𝐶𝑚𝑒𝑚)(1+𝑖𝜔𝑅𝐶ℎ𝐶𝐶ℎ) (8)

The AgPDMS electrodes will induce an electrical double layer (EDL), due to the

ionic nature of the PBS solution. In this parallel electrode configuration, the EDL is typically

modeled as an ideal capacitor Cdl, as shown in Figure 3.3 (a). The impedance of an ideal

capacitor is given by:

𝑍𝐶𝑑𝑙(𝜔) =1

𝑖𝜔𝐶𝑑𝑙 (9)

However, if the electrode is not perfectly polarizable due to inhomogeneities present

in the surface, the EDL may be better represented as a constant phase element (CPE), which

has both resistive and capacitive qualities.60, ,61 The impedance of a CPE is calculated as:

𝑍𝐶𝑃𝐸(𝜔) =1

𝑄(𝑖𝜔)𝑛 (10)

where Q represents the magnitude of the CPE, n quantifies the non-polarizable

inhomogeneities and has range 0 ≤ n ≤ 1, and ω = 2πf is the angular frequency61. It can be

seen that when n = 0, the CPE is purely resistive, and when n = 1, the CPE is purely

capacitive. The composite nature of the AgPDMS electrodes may result in the CPE as a more

accurate representation of the EDL, shown in Figure 3.3 (b). Thus, it must be determined

whether the ideal capacitor model or the CPE model more accurately represents the EDL.

The total impedance of the system is calculated as:

𝑍𝑡𝑜𝑡(𝜔) =𝑉𝑖𝑛

𝑉𝑜𝑢𝑡𝐺 (11)

where Vin = 0.4 V is the excitation voltage, Vout is the measured output of the pre-amplifier,

and G = 10 000 Ω is the gain of the pre-amplifier. Impedance data is first gathered without a

cell in the channel to fit to the appropriate model for device characterization (Rch and Cch).

Impedance data is then gathered as cells pass through the channel by the electrodes, which is

then fitted to its respective model for cell characterization (Rcyto and Cmem). The system with a

3. Fabrication and Experimental Procedure

25

cell in the channel is shown in Figure 3.3 (c) (assuming a CPE model for the EDL). From

these quantities, and using Maxwell’s mixture equation, the cytoplasm conductivity and

specific membrane capacitance are calculated as:

𝜎𝑐𝑦𝑡𝑜 =4𝜎𝑚𝑒𝑑

9𝜙ℎ𝜅𝜎𝑚𝑒𝑑𝑅𝑐𝑦𝑡𝑜−2 (12)

𝐶𝑠𝑝.𝑚𝑒𝑚 =4

9𝜙𝑟ℎ𝜅𝐶𝑚𝑒𝑚 (13)

where σmed = 1.6 S/m is the conductivity of the medium (PBS). Low frequency (11 kHz)

impedance data from 10 µm polystyrene beads (Sigma-Aldrich, Germany) are used for size

calibration, avoiding the need for high-speed imaging since the beads are well-controlled by

the manufacturer in size and shape (10.0 ± 0.2 µm).

Figure 3.3: (a) The device without a cell in the microchannel, the double layer effect is modeled as an

ideal capacitor. (b) Still without a cell, the double layer effect is modeled as a CPE. (c) The device

with a cell, which is modeled as two capacitors, representing the membrane capacitance, in series

with a resistor, representing the cytoplasm resistance.

3. Fabrication and Experimental Procedure

26

3.4 Signal Analysis

Analysis of the data is performed using MATLAB (Mathworks Inc., USA). During

experiments, raw voltage data is collected at 8 separate frequencies (11 kHz, 101 kHz, 401

kHz, 701 kHz, 1.01 MHz, 2.01 MHz, 4.01 MHz, 6.01 MHz) and converted to impedance

data using Eq. 11, as seen in Figure 3.4 (a). The data is analyzed in segments of 5 kSa in

length. The baseline is determined at each frequency by first taking a histogram of the

impedance values in the segment. A Gaussian distribution is fitted to the histogram, and the

mean of the distribution is used as the baseline impedance. Peaks are then identified using the

4.01 MHz impedance data since it has the highest SNR, as demonstrated in Figure 3.4 (b).

Corresponding peaks at other frequencies are found by taking the maximum value in a

window of 100 Sa centered around the peak sample location at 4.01 MHz. This is to account

for slight time delays in sampling at the various frequencies. The absolute impedance values

of the peaks as well as the peak heights (difference between peak impedance and baseline

impedance) are saved for further analysis.

The 11 kHz peak data is used to convert impedance to size. 10 µm polystyrene beads

are used for size calibration, which are used to identify peak data as cell events or non-cell

events. Coincident events are also omitted. Prior to experiments, PBS is loaded into the

device and a frequency sweep is performed from 1 kHz to 6 MHz to characterize the device.

The frequency sweep is curve fit to the equivalent circuit model representing the channel

without a cell, from which the device parameters are determined. These parameters are used

in the equivalent circuit model representing the channel with a cell. The absolute impedance

values of the cells are used to curve fit the appropriate circuit model, giving values for Rcyto

and Cmem, which are finally used to calculate σcyto and Csp.mem using Eq. 12 and 13,

respectively.

3. Fabrication and Experimental Procedure

27

Figure 3.4: (a) Experimental data showing peak events in a 2 second timeframe. Absolute impedance

values have been translated (and are not accurate) for the illustrative purpose of showing data from

all 8 frequencies. (b) Peak finding algorithm in MATLAB. The red line represents the baseline value,

3. Fabrication and Experimental Procedure

28

the green line represents the cut-off threshold, and the red circles highlight peaks that have been

identified.

4. Results and Analysis

29

4. Results and Analysis

4.1 AgPDMS characterization

Fabricating AgPDMS electrodes required determination of the optimal Ag concentration.

According to percolation theory, there is a threshold of conductive particle concentration at

which a continuous chain will be formed, resulting in macroscopic conductivity.47 This

threshold depends on a number of factors, such as the size of the particles, the geometry of

the particles, and the mixing procedure.47 Increasing the conductive particle concentration

beyond the threshold will yield a large increase in conductivity of the composite. The

theoretical percolation threshold was previously determined to be 16 vol % Ag.62 In practice,

AgPDMS has previously been reported to be conductive at Ag concentrations of 21 vol % (or

69 wt %) for pressure sensor applications45 and 85 wt % (or 40 vol %) for droplet

detection.57 For the sake of consistency, the lower Ag concentration of 21 vol %, though

reported as such, will be referred to as 69 wt % for the remainder of the thesis. Using a four-

point probe, we measured the conductivity of AgPDMS at different concentrations. We

found AgPDMS to be non-conductive at 69 wt % and to have excellent conductivity at 85

wt %.

Simulations were performed in COMSOL to verify conductivity on the microscale at

69 wt % Ag and 85 wt % Ag. Random distributions of Ag particles with radius 1.375 µm

were generated in a 20 by 20 µm2 square of PDMS using MATLAB and imported into

COMSOL (COMSOL, Inc., USA) (via Solidworks). Simulations were performed in 2D to

reduce computing time as a compromise for accuracy. The simulations involved the Electric

Currents physics module with one side of the square in contact with a 1 V terminal, the

opposite side of the square grounded, and the remaining two sides electrically insulated. The

conductivity and dielectric constants for PDMS and Ag were inputted as material

properties.63 Several observations were made: first, without a continuous path of Ag particles

connecting the voltage source to the ground, the conductivity is very low (~10-7 S/m),

although the 85 wt % Ag distribution had about an order of magnitude higher conductivity

4. Results and Analysis

30

than the 69 wt % Ag distribution. However, the electric field can still be seen to couple

through the Ag particles, as seen with the 69 wt % Ag distribution in Figure 4.1. Second, if

there is even only one continuous path of Ag particles from end to end (~8 vol % or 42 wt %

Ag), the conductivity of the AgPDMS will dramatically increase (~106 S/m). It can be

concluded that in practice, the percolation threshold is an empirically determined

concentration of conductive filler. Higher concentrations of conductive filler leads to a

greater probability of creating continuous end-to-end paths through the non-conductive bulk.

Figure 4.1: Electric field of 69 wt % Ag concentration simulated in COMSOL.

SEM images were taken of mock devices of 69 wt % Ag and 85 wt % Ag to observe

the Ag concentration at the sidewall. Figure 4.2 (a) and (b) show 69 wt % Ag and 85 wt %

Ag concentrations, respectively. They were taken using the backscattered electrons (BSE)

imaging mode to show the chemical composition of the bottom of the device and the sidewall

of the channel. Since heavier elements provide a higher intensity signal, the bright dots

4. Results and Analysis

31

represent Ag particles, and the remaining grey structure represents the carbon-dominant

PDMS. The density of Ag is clearly much higher in the latter case, allowing for a much

higher conductivity of the AgPDMS.

In addition, based on the equivalent circuit models regarding the modeling of the

EDL as either an ideal capacitor or a CPE, an image was taken in the standard SEM imaging

mode of the 69 wt % device to qualitatively investigate surface roughness, as seen in Figure

4.2 (c). The bumpy texture on the bottom of the device and the sidewall of the channel

caused by the Ag particles suggests the modeling of the EDL as a CPE, rather than an ideal

capacitor. Figure 4.2 (d) is a top-down view of most of the 85 wt % Ag device, showing the

overall uniformity of the AgPDMS electrodes.

4. Results and Analysis

32

Figure 4.2: SEM images showing the AgPDMS electrodes. (a) Device bottom and sidewall with 69

wt % Ag. (b) Device bottom and sidewall with 85 wt % Ag. (c) Surface roughness of bottom and

sidewall with 85 wt % Ag. (d) Top-down overview of 85 wt % Ag device.

We further optimized the conductivity of AgPDMS around the concentration of 85

wt % Ag. Thin strips of AgPDMS of various Ag concentrations were fabricated with the

same dimensions of the electrodes used in the microfluidic chip (3 mm long, 100 µm wide,

50 µm high). The conductivities are shown as a function of concentration in Figure 4.3.

While concentrations greater than 81 wt % were found to have good conductivity (> 104 S/m),

85 wt % was found to have much higher conductivity while still qualitatively having the

same ease of fabrication. AgPDMS with 87 wt % Ag was very difficult to mix, and had a

flaky texture, compared to the more paste-like composition at lower concentrations that more

closely resembled PDMS. Thus, the AgPDMS electrodes were fabricated using 85 wt % Ag

for its high conductivity, its physical consistency, and its straightforward fabrication

procedure.

4. Results and Analysis

33

Figure 4.3: Conductivity of AgPDMS with varying Ag wt %.

4.2 Device characterization

Impedance spectroscopy was performed on the microfluidic chip without cells to

determine its electrical parameters and if the CPE model was a more accurate representation

of the EDL compared to an ideal capacitor. A frequency sweep of 109 points from 1 kHz to 6

MHz was used to determine the impedance profile of the empty channel. This profile was

then fit to the empty channel equivalent circuit model using non-linear least squares curve

fitting. A sample sweep is shown in Figure 4.4, along with the fits using the CPE model for

the EDL and the ideal capacitor model for the EDL in (a) and (b), respectively. The model

using the CPE was found to have a much better fit where the average residual error at every

frequency point was on the order of 102, compared to the model using an ideal capacitor

4. Results and Analysis

34

which had an average residual error on the order of 103. Using the CPE model to represent

the EDL, the channel was found to have a resistance RCh = 23.7 ± 1.9 kΩ and a capacitance

CCh = 2.29 ± 0.18 pF. In comparison, a device with plug-in Ag/AgCl electrodes had a

resistance of RCh = 850 kΩ and a capacitance CCh = 0.18 pF. 33 The order of magnitude

difference in channel resistance allows for the sensitivity to detect particles with volume

fraction < 1%, as to be seen later. The magnitude of the CPE was found to be Q = 1.6 ± 0.2

µS with exponent n = 0.380 ± 0.007. The low uncertainty for each electrical parameter

demonstrates the electrical stability of the device. Device characterization was performed

over 3 days to investigate re-usability and possible contamination from experiments, as seen

in Figure 4.5.

Figure 4.4: Impedance profile of channels without cells along with curve fitting. (a) The fit with the

EDL represented by an ideal capacitor. (b) The fit with the EDL represented by a CPE.

4. Results and Analysis

35

Figure 4.5: Device parameters over a 3 day experiment period.

The CPE therefore exhibited characteristics closer to a resistive impedance than a

capacitive impedance due to the inhomogeneous nature of the AgPDMS composite, and does

not have as dominant an effect on the low frequency impedance data. To quantify the impact

of the EDL, we look at the ratio of the impedance of the EDL to the impedance of the

channel. For adequate sensitivity, the impedance from the channel should be greater than the

impedance of the EDL:

4. Results and Analysis

36

|𝑍𝐶ℎ(𝜔)| > |𝑍𝐶𝑃𝐸(𝜔)| (14)

|𝑅𝐶ℎ

1+𝑖𝜔𝑅𝐶ℎ𝐶𝐶ℎ| > |

1

𝑄(𝑖𝜔)𝑛| (15)

The ratio of the impedance of the channel to the impedance of the CPE at 11 kHz and

6.01 MHz is 1.3 and 13.6, respectively. Thus, the system shows adequate sensitivity for

particle detection at low frequencies. In comparison, systems that have pure metal electrodes

have been shown to have EDL with capacitances on the order of tens of pF, and thus have a

lower frequency limit closer to 500 kHz.30

4.3 Cell characterization

Impedance data was collected as cells flowed through the microchannel past the AgPDMS

electrodes. While having more frequency points is preferred, the impedance spectroscope can

at most take impedance measurements at 8 frequencies at high-speed. The frequencies were

chosen to be between 11 kHz and 6.01 MHz. Frequencies outside of this range were found to

have poor SNR due to the EDL at low frequencies and stray capacitances at high frequency.

The low end 11 kHz frequency provided size information, the high end 6.01 MHz frequency

was used to probe the cytoplasm of the cell, and the frequency points in between were used

to measure the membrane capacitance. Frequency sweeps of single cells were not possible as

with the empty channel due to practical difficulties with ‘parking’ a cell in the electrode

measurement region.

Size information is important to determine since it is required to normalize the

cytoplasm resistance Rcyto and the membrane capacitance Cmem to their intrinsic size-

independent quantities cytoplasm conductivity σcyto and specific membrane capacitance

Csp.mem, respectively. Impedance data of 10 µm beads was measured at 11 kHz for size

calibration. At low frequencies, the system is almost completely resistive, and the device

looks like a Coulter counter. As such, the impedances of beads and cells will depend only on

size, and not on capacitive properties. From 38 beads, the impedance at 11 kHz for a 10 µm

particle was found to be 19.0 ± 4.8 Ω. The distribution of impedances for the 10 µm beads is

shown in Figure 4.6.

4. Results and Analysis

37

Figure 4.6: Impedance values for 10 µm beads at 11 kHz

T24 cells were used to verify the feasibility of large data collection and cell detection.

Figure 4.7 shows a sample of 8 frequency impedance data. There is a wide range in peak size

due to other particles passing through the channel, such as PDMS debris, cell clusters, and

apoptotic cells, as observed using an optical microscope; however, based on the impedance

measurements, we were able to differentiate cell events from non-cell events. The device was

run for 40 minutes with breaks only to add more cells and PBS to the inlet. From 2400

detected peaks, we were able to identify 740 cells based on 11 kHz frequency size

information, using optical measurements and previous data as references.35 Because of the

wide channel design, the pressure was kept at 100 Pa (and not any higher) to minimize

4. Results and Analysis

38

coincident events. Some debris were determined to be over 50 µm in size; however, the wide

100 µm channel allowed passage of the debris without stopping the experiment or changing

devices. Using the 11 kHz size data, T24 was found to have an average radius of 9.1 ± 1.4

µm. In a similar manner to device characterization, the 8 frequency impedance profile was

used to curve fit the equivalent circuit model with a cell in the channel (shown in Figure 3.3

(c)). The cytoplasm conductivity and specific membrane capacitance were measured to be

0.22 ± 0.10 S/m and 30 ± 17 mF/m2, respectively. A scatter plot of the distribution of the

electrical properties is shown in Figure 4.8. Previous results found the specific membrane

capacitance to be 47.0 ± 5.1 mF/m2.35 This demonstrates the capability of the device in large

sample size enumeration and characterization while distinguishing cell events from non-cell

events.

4. Results and Analysis

39

Figure 4.7: 8 frequency impedance data with 5 particles detected in a 1 second timeframe. Based on

11 kHz peak height data, cell events were identified from non-cell events.

4. Results and Analysis

40

Figure 4.8: Scatter plot of T24 electrical properties (n = 740). Colour represents density distribution.

AML-2 and HL-60 cells were tested to measure and compare their electrical

properties. As with the T24 cells, PDMS debris, cell clusters, and apoptotic cells, as

determined by impedance amplitude data at 11 kHz, were filtered from cell events. AML-2

and HL-60 cells were found to have average radii of 8.8 ± 1.5 µm and 7.7 ± 1.4 µm,

respectively. Thus, all cell lines had a volume fraction of < 1%. The electrical properties of

AML-2 and HL-60 are compared in Figure 4.9. From 369 detected peaks, 101 AML-2 cells

were identified with a cytoplasm conductivity of 0.29 ± 0.15 S/m and a specific membrane

capacitance of 41 ± 25 mF/m2. From 733 detected peaks, 90 HL-60 cells were identified with

a cytoplasm conductivity of 0.47 ± 0.20 S/m and a specific membrane capacitance of 55 ± 26

mF/m2. These values are on the same order of magnitude as previous results and yield the

same trend that HL-60 has higher values than AML-2 (also shown in Table 3.4: AML-2 and

HL-60 electrical properties33).33,34 Differences between our values and previous results may

be due to heterogeneity and genetic drift of cell lines that result from passage number and

slight differences in culture conditions.64

4. Results and Analysis

41

Figure 4.9: AML-2 and HL-60 electrical properties. HL-60 was found to have higher cytoplasm

conductivity and specific membrane capacitance.

A potential source of error is the use of only 10 µm beads for size calibration. Typically,

beads of multiple sizes (e.g. 5 µm, 10 µm, 15 µm beads) are used to calibrate the low

frequency impedance of a particle with its size. This may provide a more accurate calibration

if the relationship between impedance and size is non-linear, which can occur at high volume

fractions.30 However, given the size of the channel in this microfluidic device (100 µm), even

particles as large as 20 µm have a volume fraction lower than 1%, which makes non-linear

effects negligible. Another possible source of error comes from the omission of certain

elements from the equivalent circuit models. Stray capacitances exist in all impedance

systems, but were not included in the equivalent circuit models since they were considered

negligible up to and including the high (6.01 MHz) frequency range. Similar devices report a

parasitic capacitance on the order of tenths of a pF.57 Adequate sensitivity for cell detection

and characterization requires that the impedance of the channel be less than the impedance of

the stray capacitance:

|𝑍𝐶ℎ(𝜔)| < |𝑍𝑠𝑡(𝜔)| (16)

4. Results and Analysis

42

|𝑅𝐶ℎ

1+𝑖𝜔𝑅𝐶ℎ𝐶𝐶ℎ| < |

1

𝑖𝜔𝐶𝑠𝑡| (17)

where Cst is the stray capacitance. From Eq. 17, the impedance of the channel at even as high

as 10 MHz is much less than the impedance of the stray capacitance. In addition, stray

capacitances tend to have a more noticeable effect in planar and plug-in electrode

configurations compared to parallel electrode configurations.28,33,65,66 This is due to coupling

of the electrodes through the substrate material instead of the intended conductive medium,

which happens to a much smaller degree when the electrodes are directly across from each

other. Thus, the stray capacitance can be ignored in our microfluidic system.

While we were able to properly identify single-cell peaks and were not hindered by

clogging of the device, the debris unnecessarily complicated analysis. PDMS is commonly

used in microfluidic applications due to its versatile properties, such as flexibility,

transparency, and biocompatibility.42 However, it has also been the source of contamination

in the form of debris, which can occur during unavoidable intubation procedures (filling PBS

into the device, loading cells at the inlet, attaching pressure source). A possible solution

would be to implement filtration structures to minimize the passage of larger debris and cell

clusters.66 This would both reduce the number of non-cell events as well as reduce coincident

events from cells and debris simultaneously passing through the channel. Overall, the device

was able to detect and electrically characterize cells with a volume fraction of < 1%.

5. Conclusion

43

5. Conclusion

This thesis reported a wide channel PDMS microfluidic chip that used AgPDMS embedded

electrodes to perform single-cell electrical characterization in the form of cytoplasm

conductivity and specific membrane capacitance. It was demonstrated that a large 100 µm

channel design could be used to extend the device lifetime by minimizing clogging issues,

allowing for repeated use over several days. Parallel AgPDMS sidewall electrodes were used

for the first time to perform impedance spectroscopy, limiting the EDL while providing the

necessary sensitivity for impedance spectroscopy. Multi-frequency impedance data was fitted

to equivalent circuit models to determine cytoplasm conductivity and specific membrane

capacitance. The T24 cell line was used to verify data measurement, with n = 740 cells being

detected. AML-2 and HL-60 cells were used to validate the characterization capabilities, with

the cytoplasm conductivities determined to be 0.29 ± 0.15 S/m and 0.47 ± 0.20 S/m, and the

specific membrane capacitances determined to be 41 ± 25 mF/m2 and 55 ± 26 mF/m2,

respectively.

5.1 Future directions

Possibilities of future works and further development of the device are as follows:

1. Addition of a second set of AgPDMS electrodes to perform a more accurate differential

measurement.

2. Characterization of rare cell types such as CTCs.

3. On-chip integration of filtration and sample processing structures. This can be used to

perform a complete blood count: red blood cell enumeration and 3-part white blood cell

differential.

4. Miniaturization of the microchannel. Assuming the same sensitivity to particles with

volume fraction < 1%, a 20 x 20 x 20 µm3 electrode measurement region could detect

platelets or even bacteria.

5. Conclusion

44

5. Integration of sorting structures to separate and collect cells.

6. Development of the model and device to characterize non-negligible membrane

resistance and cytoplasm capacitance.

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45

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Appendix

50

Appendix

A Previous electrode designs

Figure A: Previous designs were abandoned due to increased stray capacitance effect. (a) Two 100

µm branches instead of one to increasing the probability of successfully fabricating electrodes.

Second branch was found to not greatly increase success rate. (b) Thin 100 µm electrode area is

smaller. Fabrication was actually more difficult.

Appendix

51

B AgPDMS photopatterning procedure

Table B: Photopatterning procedure for 50 µm high AgPDMS

Step Procedure

1 3 wt. % benzophenone, 21 vol. % Ag added to PDMS, degassed for 15 min.

2 Pour AgPDMS + benzophenone mixture onto glass slide

3 Place slide and mask in mask aligner. Use proximity mode with 50 µm spacing.

Expose for 10 min

4 Post-exposure bake: 120 °C for 60 s

5 Develop in toluene for 5 s

6 Rinse in IPA and dry with N2 gun