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
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