석사 학위논문

Master's Thesis

면역특이적으로 크기가 증대된 혈중

희귀세포의 고순도 회수를 위한 관성

미세유체소자 내에서의 분리법

High-yield Purification of Rare Blood Cells by

Immunospecific Size Enhancement in an Inertial

Microfluidic Device

신중호 (申仲浩 Joong Ho Shin)

Department of Bio and Brain Engineering

KAIST

2014

면역특이적으로 크기가 증대된 혈중

희귀세포의 고순도 회수를 위한 관성

미세유체소자 내에서의 분리법

High-yield Purification of Rare Blood Cells by

Immunospecific Size Enhancement in an Inertial

Microfluidic Device

High-yield Purification of Rare Blood Cells by

Immunospecific Size Enhancement in an Inertial

Microfluidic Device

Advisor: Professor Je-Kyun Park

by

Joong Ho Shin

Department of Bio and Brain Engineering

KAIST

A thesis submitted to the faculty of KAIST in partial fulfillment of the re-

quirements for the degree of Master of Science in the Department of Bio and

Brain Engineering. The study was conducted in accordance with Code of Re-

search Ethics1

2013. 12. 12

Approved by

___________________

Professor Je-Kyun Park

1Declaration of Ethical Conduct in Research: I, as a graduate student of KAIST, hereby declare that I have not

committed any acts that may damage the credibility of my research. These include, but are not limited to: falsi-

fication, thesis written by someone else, distortion of research findings or plagiarism. I affirm that my thesis

contains honest conclusions based on my own careful research under the guidance of my thesis advisor.

면역특이적으로 크기가 증대된 혈중

희귀세포의 고순도 회수를 위한 관성

미세유체소자 내에서의 분리법

신 중 호

위 논문은 한국과학기술원 석사학위논문으로

학위논문심사위원회에서 심사 통과하였음.

2013 년 12 월 12 일

심사위원장

심사위원

심사위원

박 제 균 (인)

정 기 훈 (인)

박 인 규 (인)

i

MBiS

20123376

신 중 호. Joong Ho Shin. High-yield Purification of Rare Blood Cells by Immunospecific

Size Enhancement in an Inertial Microfluidic Device. 면역특이적으로 크기가 증대된

혈중 희귀세포의 고순도 회수를 위한 관성 미세유체소자 내에서의 분리법. De-

partment of Bio and Brain Engineering. 2013. 73 p. Advisor Prof. Je-Kyun Park. Text in

English

ABSTRACT

Circulating rare cells in blood provide valuable information and their separation can open opportuni-

ties for biological studies. Separation of circulating tumor cells (CTCs) in specific can be used as a non-

invasive liquid biopsy and biomarker for cancer diagnosis and screening purposes. CTCs are generally larger

than white blood cells (WBCs) and many size-based separation methods utilizing inertial microfluidics for

high throughput separation have been developed in recent years. However, due to the overlap in size distribu-

tions between WBCs and CTCs, it has been difficult to separate with both high recovery rate and purity. In

this study, to overcome the limitation of size-based inertial separation of CTCs from WBCs, immunospecific

size enhancement was applied to eliminate that size overlap. To demonstrate improved separation, microbeads

are attached to CTCs by antigen–antibody binding to enhance their sizes, and then separated from WBCs in an

inertial microfluidic separating device that consists of a series of contraction–expansion arrays (CEA). In ad-

dition, the dimensions of CEA device were systematically varied for optimal separation performance and

computational fluid dynamics simulations were performed to characterize the secondary flows in the device.

Breast cancer and monocyte cancer cell lines were chosen to model CTCs and WBCs inertial migration in the

optimized device, respectively, and then breast cancer cells were labeled with beads to compare the separation

performance. The device separates unlabeled CTCs from WBCs with 74% recovery rate and 76% WBC rejec-

tion ratio at 0.88 ml/h. However, increased recovery rate of 97.6% and 95% WBC rejection ratio at 0.88 ml/h

was demonstrated by labeling the CTCs with 5 μm beads. With increased recovery rate and purity provided by

increased WBC rejection ratio, inertial separation of bead-labeled CTCs not only has the potential to improve

the clinical diagnostic, prognostic and screening results, but also to provide invaluable opportunities for rare

cell researches, including hematopoietic stem cell, cancer metastasis, and cancer gene mutation studies.

Keywords: Inertial microfluidics, Secondary flow, Cell separation, Circulating rare cells, Circulating tumor

cells

ii

Contents

Abstract ································································································· i

Table of Contents ····················································································· ii

Nomenclature························································································· iv

List of Tables ························································································· vi

List of Figures ······················································································· vii

1. Introduction

1.1 Research Background ······································································· 1

1.2 Related Research Trends ··································································· 2

1.3 Research Purpose ············································································ 7

2. Concept and Theory

2.1 Concept ······················································································· 9

2.2 Secondary flow in a CEA microchannel ················································· 12

2.3 Inertial migration ··········································································· 15

3. Experimental

3.1 Design principle ··········································································· 17

3.2 Microchannel design and fabrication ···················································· 19

3.3 Computational fluid dynamics simulation of secondary flow ························ 22

3.4 Cell-bead labeling ·········································································· 22

3.5 Experimental setup

3.5.1 CEA device characterization with fluorescent beads ························· 24

3.5.2 Cell trajectory observation ······················································· 24

3.5.3. Image analysis ····································································· 25

4. Results and Discussion

4.1 Immunospecific size enhancement of cells ·············································· 27

4.2 Simulation of secondary flow formation ················································ 31

4.3 Contraction–expansion array number optimization ···································· 33

4.4 Contraction region length optimization ·················································· 37

4.5 Optimized channel characterization ······················································ 41

iii

4.6 Characterization of cell trajectories ······················································ 43

4.7 Cell trajectories from mixed sample ····················································· 47

4.8 Evaluation of separation efficiency ······················································ 49

5. Conclusions

5.1 Summary ····················································································· 54

5.2 Future direction ············································································· 58

6. References ·························································································· 62

Summary in Korean ················································································ 68

Acknowledgments ··················································································· 72

Curriculum Vitae ···················································································· 73

iv

Nomenclature

Alphabetic letters

Yavg Average lateral position

FD Drag force

Ρ Density of fluid

S Displacement

FL Inertial lift force

Ii Intensity value of a pixel

CL Lift coefficient

yi Normalized location of a pixel

ap Particle diameter

Re Reynolds number

Rs Separation resolution

W Smallest dimension of the microfluidic channel

σ Standard deviation

v Transverse flow component in the y-plane

w Transverse flow component in the z-plane

US Velocity of the secondary flow

μ Viscosity of fluid

Fw Wall-induced inertial lift force

Um x-axial maximum flow velocity

v

Abbreviations

BSA Bovine serum albumin

CTC Circulating tumor cell

CFD Computational fluid dynamics

CEA Contraction–expansion array

DNA Deoxyribonucleic acid

EpCAM Epithelial cell adhesion molecule

FACS Fluorescence-activated cell sorting

FDA Food and Drug Administration

fps Frames per second

HSC Hematopoietic stem cell

MHC Major histocompatibility complex

MOFF Multi-orifice flow fractionation

PBS Phosphate buffered saline

PDMS Poly(dimethylsiloxane)

PCR Polymerase chain reaction

RBC Red blood cell

rpm Revolutions per minute

SD Standard deviation

WBC White blood cell

vi

List of Tables

Table 1. Summary of current technologies utilizing inertial microfluidics for cell separation.

Table 2. Summary of average diameters and standard deviation of U937, MCF-7 cells and MCF-7 cells labeled

with 3 and 5 μm beads.

Table 3. The average and standard deviations of 4 μm particles in 4, 6, 8 array devices at a range of Re number

from 3 to 12.

Table 4. The average (AVG) and standard deviation (SD) values for U937, MCF-7 cells and 5 μm bead labeled

MCF-7 cells at different Re.

Table 5. Summary of the size-based inertial separation performance demonstrated in recent trends compared to

that achieved in this work.

vii

List of Figures

Figure 1. a) Schematic of size-based cell separation in CEA device. Sample is injected from the inlet 1 and par-

ticles are pushed to side B by focusing flow. After going through a series of contraction–expansion arrays they

are separated by size and bifurcated into two outlets. Bead-labeled CTCs are collected from the outlet 1 and

blood cells are removed through the outlet 2. b) Photo of the fabricated CEA device with orange colored dye for

visualization.

Figure 2. Schematic representation of a) overlap in size distribution of U937 and MCF-7 cells, b) separated

average lateral positions of the cells in the CEA microchannel before outlet, c) overlap in lateral position distri-

bution of U937 and MCF-7 cells, d) shifted size distribution of bead-labeled MCF-7, e) changed average lateral

position of the bead-labeled MCF-7 within the microchannel, and f) the shifted lateral position distribution of

bead-labeled MCF-7. Bifurcating the microchannel at the marked position can result in improved separation

efficiency.

Figure 3. a) Schematic of three-dimensional illustration of CEA microchannel showing locations of cross sec-

tions at which CFD analysis was performed to visualize transverse velocity components. Second expansion re-

gion was analyzed with each section being 1380, 1480, 1580, 1650 and 1730 μm away from the beginning of

the first expansion region (x = 0). b) Projections of v components on the planes perpendicular to x-axis. The 5th

plane shows the counter rotating vortices, which indicates the formation of secondary flow at the entrance of the

contraction region.

viii

Figure 4. Magnitude of v components at each sections of expansion region. The magnitude of the transverse

components are larger near the entrance of the contraction region.

Figure 5. Schematics of a) particles with a diameter a flowing through microchannel reaching equilibrium posi-

tions due to the balance between wall-induced (FW) and shear-induced lift forces (FL) and b) randomly dis-

persed particles with distinct equilibrium positions in square and low aspect ratio (ratio of height to width is less

than 1) channel.

Figure 6. Schematic of a) CEA microchannel with entrance, contraction region and exit indicated and b) the

lateral positions (with respect to y axis) of relatively larger and smaller particles as at each section. The distance

between two different sized particles, indicated by ΔS, increases as they flow through more arrays.

Figure 7. Illustration of photolithography, PDMS mold casting and glass bonding for microchip fabrication.

Figure 8. Dimensions of the CEA microchannel design. Expansion region was expanded before bifurcation

point to observe cell’s trajectories at higher resolution.

Figure 9. Schematic of labeling MCF-7 cell with a microbead. Strepdavidin coated microbead and biotinylated

antibody are bound first, then MCF-7 cell is labeled with the bead by antibody and EpCAM binding.

Figure 10. Picture of the setup for separation experiment. Sample is injected from the inlet 1 and buffer is in-

ix

jected from the inlet 2 with syringe pumps and the lateral positions of the cells can be monitored.

Figure 11. a) Size distribution of U937, MCF-7 cells and MCF-7 cells labeled with 3 and 5 μm beads. Labeling

MCF-7 cells with 5 μm beads completely overcomes the overlap in size distribution with U937 cells. Pictures of

b) poorly labeled MCF-7 cell, c) bead-covered MCF-7 cell showing uncovered spot, and d) well covered MCF-

7 cell. All pictures were taken from same incubation batch to show that there is variation in coverage.

Figure 12. Simulation results of a) vw components and b) curl field projected on the second contraction region,

30 μm inward the entrance and c) the magnitude of vw and curl values plotted. The three dimensional schematic

at the top of the figure illustrates the cross sectional plane (red dotted square) at which simulations were ob-

served.

Figure 13. Schematic of particle’s lateral position divergence due to recirculation of secondary flow. Particle’s

lateral positions, after being focused at side A, can diverge after passing through more arrays by entraining in

the secondary flow and recirculating back to side B.

Figure 14. Streamlines formed by 4 μm diameter microbeads captured at the end the of a CEA channel with a)

4 arrays, b) 6 arrays, and c) 8 arrays.

Figure 15. Plot of the normalized average lateral positions of 4 μm particles in CEA device with 4, 6, and 8

arrays.

x

Figure 16. The trajectories of a) 15 μm and b) 25 μm diameter beads observed in CEA devices having 150, 300

and 450 μm long contraction regions. In each devices, the trajectories were observed at a range of Re values

from 3 to 12. Scale bar: 200 μm

Figure 17. Plot showing the normalized lateral positions of 15 and 25 μm fluorescent microbeads in a CEA de-

vice with a) 150 μm, b) 300 μm, and c) 450 μm long contraction region lengths for Re from 3 to 12.

Figure 18. Graph of a) the normalized lateral positions of 4, 15 and 25 μm beads plotted against Re, and b) the

normalized lateral positions versus particle diameter.

Figure 19. Image of a) U937 cells, b) MCF-7 cells and c) MCF-7 cells labeled with 5 μm captured at a wide

section of CEA microchannel after passing through 4 arrays. All images are captured at Re = 12.

Figure 20. The lateral positions of U937, MCF-7 cells versus Re and the lateral positions of MCF-7 cells la-

beled with 5 μm at Re values of 9 and 12.

Figure 21. The distributions of U937, MCF-7 cells and 5 μm cell–bead complexes at Re 9 and 12.

Figure 22. a) Image of U937 cells (white sphere), unbound 5 μm streptavidin coated beads (black spots) and

MCF-7 cells labeled with 5 μm diameter beads (black sphere) flowing at Re = 12. b) The distribution of lateral

positions of U937 cells and bead-labeled MCF-7 cells.

xi

Figure 23. Illustration of a) U937 cells not being rejected by placing bifurcation point too close to side A, b)

bead-labeled MCF-7 cells being collected from outlet 1 and U937 cells being removed from the outlet 2 by ide-

al bifurcation placement, and c) bead-labeled MCF-7 cells being removed, resulting in cell loss by placing the

bifurcation too close to side B.

Figure 24. Image of cells flowing through the wide section of the CEA microchannel imaged at Re = 12. a)

U937 cells flowing and reaching the bifurcation near side A (outlet 2). b) Bead-labeled MCF-7 cells flowing and

reaching the bifurcation near side B (outlet 1). Scale bar: 300 μm.

Figure 25. MCF-7 recovery rate and U937 cell rejection ratio curve of a) unlabeled MCF-7 cell separation at Re

= 12 and b) bead-labeled MCF-7 separation at Re = 12.

- 1 -

Chapter 1. Introduction

1.1 Research Background

Separation of specific kinds of cells from other cells has been an important part of biological and bi-

omedical research. For example, staining and analyzing lymphoma cells can provide insights to rare genetic

variants of immunoglobulin or major histocompatibility complex (MHC) gene expression [1]. Isolation of hem-

atopoietic stem cell (HSC) is also important for their ability to regenerate all of the hematopoietic lineages in

vivo [2], and because it is the only stem cell that is clinically applied to treat diseases such as leukemia and con-

genital immunodeficiency [3]. Such cells however, often exist as rarely as one in a few millions of white blood

cells (WBCs) and about a billion of red blood cells (RBCs). Due to their rare events in a large population, they

are called rare cells.

Separation of rare cells is not only important for research, but it also provides valuable information

regarding a person’s health or course of a certain disease for a patient. For example, another kind of rare cell of

great interest is the circulating tumor cell (CTC), which circulates in the blood stream after being detached from

the primary tumor. CTCs have the potential to be used as screening, diagnostic and prognostic marker in cancer

patients. Isolation of CTCs is regarded as a noninvasive “liquid biopsy” for therapy guidance; and enables per-

sonalized treatment by molecular profiling [4].

Due to the extremely low number of rare cells, it is necessary to process large amount of sample in

order to collect enough cells that can provide meaningful data [1], which calls for the need of high throughput

method with high recovery. High recovery is also important for circulating tumor cell detection because it is

directly related to a cancer patient’s diagnosis and it reduces the chance of false negative test results. The col-

lected sample must also be free from presence of other blood cells for accurate molecular analysis. Several tools

- 2 -

have been utilized extensively throughout the history of biological and biomedical research, however, they are

expensive, complicated and require maintenance and skilled personnel to operate. Thus innovative methods

utilizing microfluidics have been developed to overcome the disadvantages of conventional methods.

1.2 Related Research Trends

Conventional cell separation tool in biomedical research that is widely used today is fluorescence-

activated cell sorting (FACS) [5]. Despite its high throughput and sorting efficiency, the device is expensive,

mechanically complex and require skilled technician to operate [6]. For cases of detecting circulating tumor

cells (CTCs), Veridex’s CellSearch is widely used as a current gold standard method that is FDA approved [7].

CTCs of epithelial origin express epithelial cell adhesion molecules (EpCAM) on the cell surface. CellSearch

first immunomagnetically label and separate the cells then stain the separated cells for analysis. However, the

device is also expensive and is semi-automated, which means it requires trained personnel to operate. To over-

come such limitations of conventional sorting machines, many novel microfluidic cell sorting systems have

been developed.

For separation of CTCs, the difference in size and deformability [8-10], dielectric properties [11, 12],

magnetic labeling [13] and surface expressions [14-16] between CTCs and were exploited to separate CTCs

from blood in microfluidic systems. The average sizes of CTCs are generally known to be larger and their stiff-

ness is also greater compared to those of WBCs. Thus, using filters and traps to separate CTCs by size and de-

formability can be advantageous for separating CTCs without known surface biomarkers. However, the method

can easily lose small CTCs whose sizes are smaller than the filter gaps. The devices can also result in low sepa-

ration efficiency at high flow rates due to the CTCs escaping out of the traps, which also limits the throughput

- 3 -

of optimal separation. Dielectric force based separation systems show high separation efficiency, however, the

purity decreases with increasing flow rate. The force that deviates flow path of cells is applied for shorter period

of time at higher flow rates and the cells do not separate as much as they would at lower flow rate, which gives

the system a limitation in throughput. Circulating cells that express EpCAM on the surface can be captured by

using microfluidic devices that have anti-EpCAM antibodies coated on the surface of the microchannel. Using

such method can capture CTCs with high efficiency but the purity is low and it is difficult to find the captured

cells in wide active capture area, which makes it difficult for enumeration and characterization. Separation

methods involving magnetic labeling utilizes strong magnetic attraction force to capture the labeled cells, how-

ever, cell loss can occur during washing process of WBCs and during retrieval process.

Recent advances in inertial microfluidics utilizing a secondary flow and inertial migration phenome-

non paved the way for continuous size-based separation with high throughput [17, 18]. Drag force that occur

from the formation of secondary flow and the inertial lift force involved in inertial migration become more ef-

fective at a higher flow rate. Thus by controlling these forces, particles can be ordered [17], separated by sizes

[19] and their concentrations can be enriched [20] at high throughput without the use of any external forces.

Several groups used one of the forces or combination of the two forces to separate CTCs from human blood as

summarized in Table 1 [21-23].

Separation performance is mainly judged by three criteria: CTC recovery, purity and throughput.

CTC recovery is defined as the ratio of the number of collected CTCs to the number of injected CTCs. Purity is

defined as the ratio of the number of collected CTCs to total number of cells collected. It is often represented by

WBC rejection ratio. Throughput is the volume of sample that can be processed in a given time. CTC separation

demonstrated in multi-orifice flow fractionation (MOFF) utilizes inertial lift force induced lateral position dis-

- 4 -

placement and performs at a high recovery rate of 93.75% and 90.8% WBC rejection ratio [21]. 90.8% rejection

ratio in actually means that collected sample contains WBCs in the order of 106 cells, which significantly reduc-

es the sample purity. Centrifuge-on-a-Chip technology that utilizes inertial lift force to selectively trap CTCs in

microvortices perform CTC separation with 100% purity [23]. However its recovery rate is 25.8% and such cell

loss can be detrimental to clinical applications. CTC separation performed by spiral microchannel performs at a

recovery rate of 85%, however its purity is near 5%. Such low purity can make the cancer cell detection process

difficult and increase analysis time.

Although inertial microfluidics enables high throughput separation, no device has been able to fulfill

both high recovery and purity at the same time. This limitation of size-based inertial separation stems from the

size distribution of the CTCs and WBC population. Although the average size of CTCs is known to be larger

than WBCs, there is an overlap in the size distribution of the two populations. This overlap determines the cut-

off threshold size value at which cells are to be separated. If the cutoff size is based on the largest size of WBC,

the device will remove all particles below that size, which includes all WBC population and some overlapping

smaller CTCs. The collected sample will theoretically have 100% CTC purity but it will have low recovery. On

the other hand, if the threshold size is based on the smallest size of CTC, the device will collect all CTCs, in-

cluding some of the bigger WBCs whose sizes overlap with that of CTCs. The collected sample will theoretical-

ly have 100% CTC recovery but have low purity due to the presence of WBCs.

In order to overcome the limitation caused by size-overlap, Kim et al. attached microbead on CTCs

by immunoaffinity [24, 25] to amplify their sizes, and then separated the microbead-attached CTCs from WBCs

by using filtration method with high separation efficiency. However, during the collection process of the trapped

cells through the inlet, WBCs can also be collected and can affect the molecular and genomic profiling results.

- 5 -

Furthermore, the recovery rate of the trapping method decreases with increasing flow rate, which is a major

disadvantage when considering the throughput. Thus, a new method has to be developed to overcome the innate

size-distribution overlap of sample and perform high throughput cell separation with both high recovery and

purity.

- 6 -

Table 1. Summary of current technologies utilizing inertial microfluidics for cell separation.

Hyun, Biosens. Bioelec-

tron. (2013) [21]

Sollier, Lab Chip (2014)

[23]

Warkiani, Lab Chip

(2014) [22]

Schematics

CTC Recovery (%) 93.75 25.8 85

Purity N/A 100 5

WBC Rejection (%) 90.8 N/A N/A

Throughput (ml/h) 0.6 240 102

Utilized forces Inertial lift force Inertial lift force Inertial lift force and

Dean force

- 7 -

1.3 Research Purpose

In this study, high throughput continuous separation of size-enhanced CTCs was demonstrated with

high purity and recovery rate by using an inertial microfluidic separation device. MCF-7 cells were chosen to

model CTCs that express EpCAM. U937 monocyte cell line was used to model the subpopulation of white

blood cells that are larger than the rest of white blood cells. The cell’s size distribution overlaps with that of

MCF-7 cells so that the experimental separation result does not overestimate the actual separation efficiency of

actual blood sample, which would mainly consist of much smaller white blood cells. MCF-7 cells were labeled

with 5 μm beads and their sizes were measured to show that the labeling effectively eliminates the overlap in

size distribution. Bead-labeled cells’ surfaces were analyzed to measure how much of the cell’s area was cov-

ered with beads.

The bead-labeled MCF-7 cells were injected into a contraction–expansion array (CEA) microchannel,

an inertial microfluidic device which has shown various applications, including three-dimensional focusing [26],

rapid fluid mixing [27, 28], size-based separations such as blood plasma separation [29, 30], particle separation

[31] and CTC separation from whole blood [32]. Although the device has shown high throughput separation of

CTCs with high recovery rate, its rejection ratio and purity still needs to be improved, thus the efficacy of im-

proved separation by bead-labeling was demonstrated in an optimized CEA microchannel device. In order to

determine an optimal condition, fluorescent beads were used to test various numbers of array and various con-

traction region lengths for optimal cell separation. Computational fluid dynamic simulations were also per-

formed to characterize the formation of secondary flows in the microchannel.

In a CEA device with optimized design, U937, MCF-7 cells and bead-labeled MCF-7 cells were in-

jected to characterize the lateral positions of different sized cells at various flow rates. Finally bead-labeled

- 8 -

MCF-7 cells were separated from a mixture of solution containing U937 cells and bead-labeled MCF-7 cells

and their recovery rate and WBC rejection ratio were calculated. The calculated values were then compared to

those of unlabeled MCF-7 cell separation to prove that bead-labeling results in increased separation efficiency.

- 9 -

Chapter 2. Concepts and Theory

2.1 Concepts

In this paper we demonstrate a method to improve separation efficiency of size-based inertial separa-

tion in a contraction–expansion array (CEA) microchannel by size enhancement of specific cells of interest

through immunoaffinity between cells and antibody-coated microbeads. We demonstrate the method by separat-

ing MCF-7 cells from a cell mixture containing U937 cells. The device is consisted of poly(dimethylsiloxane)

(PDMS) based microfluidic channel that has a series of wide and narrow regions and is sealed with glass slide.

This device has two inlets; one for sample injection and the other for buffer injection as shown in Figure 1. The

device also has two outlets, one for cells of interest and the other for waste. As particles flow through the micro-

fluidic channel they are exposed to inertial lift force and drag force, whose magnitudes depend on the size of the

particle. The two forces act in opposite direction to cause lateral displacement of particle across the width of the

channel, and the accumulation of the displacement results in size-based separation.

Although the difference in the average size between U937 (13.29 ± 1.40 μm) and MCF-7 cells (17.65

± 3.11 μm) (Figure 2a) result in separated average streamlines (Figure 2b) at the end of the CEA microchannel,

the frequency distribution of the lateral positions of the cells overlap each other. This overlap in lateral positions

results in reduced separation performance. However, labeling MCF-7 cells with beads shifts the size distribution

away from that of U937 cells (the actual difference between the average size becomes 14.84 μm) as shown in

Figure 2d and overcomes the overlap in size distribution. The resulting average lateral position of the bead-

labeled MCF-7 is also farther away from that of U937 cells (Figure 2e), and the overlap in the lateral position

distribution is also eliminated as a result. Finally, by bifurcating the channel in half with two outlets, the bead-

labeled MCF-7 cells can be collected from the outlet 1 (from side B) improved purity and recovery.

- 10 -

Figure 1. a) Schematic of size-based cell separation in a contraction–expansion array (CEA) device. Sample is

injected from the inlet 1 and particles are pushed to side B by focusing flow. After going through a series of

contraction–expansion arrays they are separated by size and bifurcated into two outlets. Bead-labeled CTCs are

collected from the outlet 1 and blood cells are removed through the outlet 2. b) Photo of the fabricated CEA

device with orange colored dye for visualization. Scale bar = 5 mm.

- 11 -

Figure 2. Schematic representation of a) overlap in size distribution of U937 and MCF-7 cells, b) separated

average lateral positions of the cells in the CEA microchannel before outlet, c) overlap in lateral position distri-

bution of U937 and MCF-7 cells, d) shifted size distribution of bead-labeled MCF-7, e) changed average lateral

position of the bead-labeled MCF-7 within the microchannel, and f) the shifted lateral position distribution of

bead-labeled MCF-7. Bifurcating the microchannel at the marked position can result in improved separation

efficiency.

- 12 -

2.2 Secondary flow in a CEA microchannel

Fluid flows slower in a channel with wider cross-sectional area and faster in narrower cross-sectional

area to maintain a constant volumetric flow rate. In CEA microchannel, wide section of the fluid channel be-

comes narrow along one side of the channel (side A) as shown in Figure 3a. Figure 3b shows that fluid near side

B accelerates towards side A to enter the narrow region and the resulting flow profile obtains V component

(along the y-axis), whose direction is orthogonal to the direction of main flow (along the x-axis). The magnitude

of the transverse flow increases as fluid nears the entrance of contraction region as shown in the graph (section

1 to 4). The fluid that accelerated toward side A diverges to both side of z-axis and recirculates back toward side

B (section 5), thus forming secondary flow at the entrance. Figure 4 shows the velocity profile of the secondary

flow composed of two symmetric, counter rotating vortices, which is similar to Dean vortices observed in

curved channel [33].

Assuming stokes drag, particles flowing through the entrance experience drag force defined as below

[31] :

psD 3 aUF (1)

where μ is the viscosity of the fluid, US is the velocity of the secondary flow (transverse velocity) and ap is the

particle diameter. The force, which depends on the diameter of the particle and can be controlled by changing

the transverse velocity, deflects particle’s trajectory to side B upon its entrance to contraction region.

- 13 -

Figure 3. a) Schematic of three-dimensional illustration of CEA microchannel showing locations of cross sec-

tions at which CFD analysis was performed to visualize transverse velocity components. Second expansion re-

gion was analyzed with each section being 1380, 1480, 1580, 1650 and 1730 μm away from the beginning of

the first expansion region (x = 0). b) Projections of V components on the planes perpendicular to x-axis. The 5th

plane shows the counter rotating vortices, which indicates the formation of secondary flow at the entrance of the

contraction region.

- 14 -

Figure 4. Magnitude of v components at each sections of expansion region. The magnitude of the transverse

components are larger near the entrance of the contraction region.

- 15 -

2.3 Inertial migration

Neutrally buoyant particles flowing in a straight channel experience inertial migration from the center

toward the wall of the channel and reach equilibrium positions. Segre and Silberberg were the first to report

particles equilibrating at ~0.2D away from a cylindrical wall with channel diameter D [34]. Similar phenome-

non has been observed in microchannels of square and rectangular cross sections with distinct equilibrium posi-

tions [17, 35, 36]. The migration of particles is caused by the parabolic velocity profile of Poiseuille flow [37]

and the force that causes particles to migrate across the streamline toward channel wall is related to the shear

rate across the channel wall (Figure 5a). The force, which is known as shear-induced lift force, is defined as

below [37, 38]:

2

L

4

p

2

m

LW

CaUF

(2)

where ρ, Um, ap, CL, and W are density of the fluid, x-axial maximum flow velocity, diameter of the particle, lift

coefficient, and the smallest dimension of the microfluidic channel, respectively.

As a particle nears the channel wall, it is pushed away from the wall due to asymmetrical wake

around the particle, and this force is defined as wall-induced lift force (Fw) [39]. The particle flowing through

the channel reaches one of multiple equilibrium positions where the two lift forces are balanced. The number of

equilibrium positions depends on the aspect ratio of the channel: four for square channel and two for rectangular

channel as shown in Figure 5b. Particles equilibrate at the center of each side wall for square channel and they

equilibrate at the center of longer walls for rectangular channel [18].

Inertial lift force mainly affects particle’s trajectories in the contraction region of the CEA micro-

channel. Because the particles are focused with focusing flow and enter the contraction region at the wall of

side B initially, the lift force acts only toward side B until the particles cross the centerline of the channel. Due

- 16 -

to the ap4 term in equation (2), larger particles migrate back to the equilibrium position (toward side B) faster

than smaller particles do after being deflected by the secondary flow.

Figure 5. Schematics of a) particles with a diameter a flowing through microchannel reaching equilibrium posi-

tions due to the balance between wall-induced (FW) and shear-induced lift forces (FL) and b) randomly dis-

persed particles with distinct equilibrium positions in square and low aspect ratio (ratio of height to width is less

than 1) channel.

- 17 -

Chapter 3. Experimental

3.1 Design principle

The CEA microchannel utilizes two forces to separate particles based on their size: drag force and in-

ertial lift force. Particles in the sample are initially pushed to side B by the buffer flow as indicated in Figure 6a.

The formation of secondary flow at the entrance of contraction region causes deflection in particle’s trajectory

toward side A. Particles are then exposed to inertial lift force and migrate back to side B due to inertial lift force.

As shown in equation (2), the lift force scales as FL ~ ap4. Thus larger particle (orange) migrates back to the

equilibrium position (toward side B) much faster than smaller (green) particle does. This causes separation

(ΔS1) observed at the exit of contraction region. The particles maintain their streamline until they enter the sub-

sequent contraction region and the deflection in trajectory caused by the two forces repeated, which results in

farther separation between the two particles (ΔS2). As particles pass through more arrays, they result in farther

separation from each other (ΔSn).

The extent to which particles migrate back to side B is proportional to the length of the contraction

region, which is also proportional to the period of time at which inertial lift force is applied. The length of the

contraction region can be controlled so that particles that are larger than certain threshold size can migrate back

to side B while particles below the size threshold can migrate toward side A. Flowing through additional arrays

eventually leads to farther spatial separation between different sized particles. However, because the secondary

flow occurs in a circular manner, the particles that have migrated to side A of the channel may migrate back to

side B upon the entrance of additional array as shown in Figure 13. The particle stream that was focused at side

B entrains in the secondary flow and diverges as the flow recirculates back to side B. Thus proper selection of

array number is required for focusing of smaller particles.

- 18 -

Figure 6. Schematic of a) CEA microchannel with entrance, contraction region and exit indicated and b) the

lateral positions (with respect to y axis) of relatively larger and smaller particles as at each section. The distance

between two different sized particles, indicated by ΔS, increases as they flow through more arrays.

- 19 -

3.2 Microchannel design and fabrication

The CEA microchannel was fabricated by soft lithography techniques as illustrated in Figure 7. SU-8

negative photoresist was spin-coated on silicon wafer. A film mask with microchannel design was placed on top

and microstructures were patterned by ultraviolet light exposure. PDMS prepolymer was mixed with its curing

agent (Sylgard 184; Dow Corning, MI) in the ratio of 9:1, poured on the SU-8 photoresist mold, and then cured

for 30 min on a hot plate at 85 ˚C. The PDMS replica was irreversibly bonded with glass slide by oxygen plas-

ma treatment. The design of CEA microchannel is shown in Figure 8.

- 20 -

Figure 7. Illustration of photolithography, PDMS mold casting and glass bonding for microchip fabrication.

- 21 -

Figure 8. Dimensions of the CEA microchannel design. Expansion region was expanded before bifurcation

point to observe cell’s trajectories at higher resolution.

- 22 -

3.3 Computational fluid dynamics simulation of secondary flow

Computer fluid dynamics (CFD) simulation was performed to observe the formation of secondary

flow at different flow rates. CEA microchannel channel that consists of 6 arrays with 300 μm long contraction

region with 55 μm height was constructed using a commercial geometry modeler (CFD-GEOM; ESI Group,

Huntsville, AL) and multiphysics solver (CFD-ACE+; ESI Group) was used to simulate the flow characteristics.

The density of water was set to 997 kg m−3

and dynamic viscosity was set to 8.55 × 10−4

kg m−1

s−1

. The applied

total flow rates were varied from 2.19 to 8.75 mL h−1

, while the outlet was set to a fixed-pressure boundary con-

dition.

3.4 Cell–bead labeling

To overcome the size-overlap between rare cells and WBCs, polystyrene beads were attached to

MCF-7 cells via immunoaffinity as shown in Figure 9. Beads are coated with antibodies that bind to EpCAM

first so that they can bind to the proteins expressed on the cell’s surface. 5 μm diameter streptavidin coated

beads (Bangslab, Fishers, IN) were conjugated with monoclonal antihuman EpCAM/TROP1 biotinylated anti-

body (R&D Systems, Minneapolis, MN). Beads and antibodies were rotated at 10 revolutions per minute (rpm)

on a rotator (PTR-60, Grant-bio, Shepreth, U.K) for 1 h while being suspended in phosphate buffered saline

containing (PBS) 1% bovine serum albumin (BSA) for streptavidin-biotin interaction. Beads were washed three

times to remove any unbound antibody, and rotated again at 10 rpm with MCF-7 cells in PBS containing 1%

BSA at room temperature for antibody–antigen binding in 5.73 × 107 beads.ml concentration to ensure cell–

bead collision and thorough labeling [24]. For characterization of bead-labeled cell’s trajectories by themselves,

MCF-7 cells were incubated with the beads for 1 h. For separation experiment that was performed with U937

- 23 -

cells, MCF-7 cells were incubated with the beads for 2 h. Microscopic images of the bead-labeled cells were

taken while the view was focused at the top of the cell (away from the glass slide along the z-axis) so that the

upper part of the cell’s surface can be imaged. The coverage of cell–bead labeling is defined as the ratio of sur-

face area occupied with beads to the total cross sectional area of cell–bead complex.

Figure 9. Schematic of labeling MCF-7 cell with a microbead. Strepdavidin coated microbead and biotinylated

antibody are bound first, then MCF-7 cell is labeled with the bead by antibody and EpCAM binding. Note that

image is not scaled proportionally.

- 24 -

3.5 Experimental setup

3.5.1 CEA device characterization with fluorescent beads

4 μm diameter red fluorescent sulfate beads, 15 μm diameter red fluorescent polystyrene beads and

25 μm diameter green fluorescent latex beads were used as models of red blood cells (RBCs), white blood cells

(WBCs) and bead attached MCF-7 cells respectively to characterize the CEA device. 4, 15 and 25 μm diameter

beads were each suspended in 0.2 % Pluronic solution (Sigma-Aldrich Co., St. Louis, MO) with concentration

of 2 × 106, 5 × 10

5, and 5 × 10

4 beads/ml respectively. 0.2% Pluronic solution was used to prevent the beads

from adhering to the wall and to each other, and it was also used as a focusing flow. The beads were injected as

a sample flow from inlet 1 and focusing flow was injected from inlet 2 with a ratio of 1:9. The total flow rates

tested are 2.19, 4.37, 6.56 and 8.75 ml/h, which are equivalent to Reynolds (Re) number of 3, 6, 9 and 12 when

converted with respect to the expansion region dimension. Injection was performed with syringe pumps

(KDS200; KD Scientific Inc., Holliston, MA). The trajectories of the fluorescent beads were imaged using a

fluorescence microscope (TS100; Nikon Co., Japan) with a charge-coupled device (DS-2Mv; Nikon Co.) with

an exposure time of 1 s.

3.5.2 Cell trajectory observation

U937 cells and MCF-7 cells were used as models of WBC and CTC. U937, MCF-7 and bead-labeled

MCF-7 cells were injected separately at concentration of 1.25 × 106, 5 × 10

5, and 10

4 cells/ml to observe their

respective lateral position within the channel. 2 × 104 cell–bead complex composed of bead-covered MCF-7

cells were spiked into 2 ml of 1.25 × 106 cells/ml U937 suspension to observe the effect of WBC’s presence on

inertial migration of cell–bead complex. PBS was used as a focusing flow for all cell trajectory experiments and

- 25 -

the total flow rates tested were the same as in device characterization with beads. Cells were imaged by high

speed camera (HotShot 512 sc; NAC, Simi Valley, CA) at 10,000 frames per second (fps). Picture of overall

setup, including microfluidic device, pumps, syringes, and camera, is show in Figure 10.

Figure 10. Picture of the setup for separation experiment. Sample is injected from the inlet 1 and buffer is in-

jected from the inlet 2 with syringe pumps and the lateral positions of the cells can be monitored.

3.5.3. Image analysis

The sizes of cells and cell–bead complexes were measured by injecting them separately in disposable

hemocytometer (DHC-N01; INCYTO, Korea) and analyzing the bright field images with ImageJ software. The

counting grids of disposable hemocytometers were used as reference dimension for calibration. The images of

trajectory bands formed by fluorescent beads were stacked and the fluorescent intensity profile across the ex-

pansion region was obtained using ImageJ software. The average lateral position Yavg of the beads was calculat-

ed according to the following equation:

- 26 -

n

i

n

i

I

IyY

1 i

1 ii

avg (3)

where yi and Ii are normalized location of pixels relative to the width of the channel and intensity value at each

pixel, respectively. Standard deviation of fluorescent beads is defined as:

n

i

n

i

I

IYy

1 i

1 i

2

i )( (4)

which is used to calculate the separation resolution between the average lateral position of two different diame-

ter particles. The separation resolution is defined as follows:

)(5.021

s

YR (5)

Where ΔY is the difference between the average lateral position of two particles and σ is the standard deviation

of lateral position of each particle. The lateral positions of cells with respect to the width of the expansion re-

gion were measured using ImageJ software with side A as the baseline of lateral position.

- 27 -

Chapter 4. Results and Discussion

4.1. Immunospecific size enhancement of cells

In order to demonstrate that our approach can effectively separate specific cell type from others de-

spite their size overlap, two types of cell lines with overlapping size distribution were chosen: MCF-7, a breast

cancer cell line; and U937, a monocyte cancer cell line. U937 cells were chosen due to their relatively larger

size compared to other WBC types, whose size distribution overlaps with that of MCF-7 cells. As shown in Fig-

ure 11, U937 cells (average size = 13.29 ± 1.40 μm) and MCF-7 cells’ (average size = 17.65 ± 3.11 μm) size

distribution overlaps. To overcome the size overlap by labeling MCF-7 cells with beads, 3 and 5 μm beads were

attached and the enhanced sizes were measured. As it can be shown from the plot attachment of 3 and 5 μm

beads to MCF-7 results in increased diameter of 24.09 ± 3.90 and 28.13 ± 3.17 μm, respectively, and shifts the

size distribution of cell–bead complex away from U937’s size distribution. Attaching 3 μm diameter beads in-

creases the average diameter of MCF-7 cell by 6.44 μm, however, does not completely remove the overlap in

size distribution. This is due to the variation in MCF-7 cell’s size, whose standard deviation is 3.11 μm. Thus we

chose 5 μm bead for size-amplification to increase the average size by 10.48 μm and eliminated the size overlap.

The resulting lateral position difference caused by increased effect of inertial lift force that arises from size dif-

ference would also theoretically be greater compared to that caused by size-amplification with 3 μm beads by a

factor of 1.85 according to equation 2. The coverage obtained with 5 μm bead labeling was calculated to be

90.97% (n = 18), which suggests that the variation in the increased amount of inertial lift force by the size en-

hancement is inconsiderable due to uniformly increased size.

As mentioned previously, the coverage seen from the bottom and side views of the bead-labeled cells

are not considered for the coverage calculation and the top view is assumed to represent the overall coverage.

- 28 -

This assumption is not an accurate way to determine the coverage, however, it does not overestimate the overall

coverage because more beads are theoretically attached at the bottom compared to the top. As the suspension

containing bead-labeled cells are dispensed on the glass slide for microscopic imaging, the side that has more

beads attached would turn toward the bottom due to the gravity, consequently putting the less covered side to

the top, and thus the top part was imaged for conservative estimation of the coverage. Although the extent to

which beads label MCF-7 cells varies as shown in Figure 11b-d, no MCF-7 cells were observed without any

beads attached to them. Imperfect coverage may be due to not enough incubation time for cell to bead interac-

tion and also due to a sphere packing issue [24]. The number of binding sites is not regarded as a limiting factor

in cell coverage considering the high level of EpCAM expression (2 × 105 binding sites per cell) [40].

- 29 -

Figure 11. a) Size distribution of U937, MCF-7 cells and MCF-7 cells labeled with 3 and 5 μm beads. Labeling

MCF-7 cells with 5 μm beads completely overcomes the overlap in size distribution with U937 cells. Pictures of

b) poorly labeled MCF-7 cell, c) bead-covered MCF-7 cell showing uncovered spot, and d) well covered MCF-

7 cell. All pictures were taken from the same incubation batch to show that there is variation in coverage.

- 30 -

Table 2. Summary of average diameters and standard deviation of U937, MCF-7 cells and MCF-7 cells labeled

with 3 and 5 μm beads.

U937

(n = 60)

MCF-7

(n = 45)

MCF-7 with 3 μm

bead

(n = 22)

MCF-7 with 5 μm

bead

(n = 22)

Diameter (μm) 13.29 ± 1.40 17.65 ± 3.11 24.09 ± 3.90 28.13 ± 3.17

- 31 -

4.2. Simulation of secondary flow formation

Computational fluid dynamics simulation was performed to observe the formation of secondary flow.

CEA device with 6 arrays was simulated to obtain the magnitude of the transverse flow and the respective curl

values at the 2nd entrance (30 μm inward) of contraction region. Scalar projection of velocity vectors on cross

section (vw components on a plane orthogonal to x-axis) of contraction region was obtained to visualize the

transverse vector components that are orthogonal to the main flow direction (x-direction) and Figure 12a shows

the projected vector field in a form of two counter rotating vortices. The interface of the two vortices is at the

middle of the channel and directs toward side A, which causes particles to migrate side A as they enter the con-

traction region. The magnitude of vw components were measured for Re of 3 to 12 to observe the change in

magnitude with respect to increasing flow rate. Scalar values of curl of the velocity field at the cross section are

also investigated as shown in Figure 12b. Scalar values of curl vectors along the x axis were obtained to visual-

ize the magnitude of rotation in transverse direction. The magnitude of the maximum vw component at 30 μm

inward from the entrance was measured from Re of 3 to 12 and their respective maximum curl magnitudes were

also measured and are plotted in Figure 12c. The plot shows that the maximum vw component increases with

increasing Re, which indicates that the magnitude of drag force on particles caused by secondary flow becomes

stronger. The magnitude of curl also increases with increasing Re.

- 32 -

Figure 12. a) The three dimensional schematic at the top of the figure illustrates the cross sectional plane (see

red dotted square) at which simulations were observed. Simulation results of b) vw components and b) curl field

projected on the second contraction region, 30 μm inward the entrance and d) the magnitude of vw and curl val-

ues plotted.

- 33 -

4.3. Contraction–expansion array number optimization

The overall goal of the device is to remove RBCs, WBCs and bead-labeled CTCs from whole blood

simultaneously. As mentioned previously, after the deflection in lateral positions of particles toward side A upon

the entrance of contraction region, larger cells migrate back to side B quicker than smaller cells (RBCs) do.

Thus, increasing the number of arrays mainly affects the lateral position of RBCs and adding too many arrays

may diverge the lateral position of the cells due to recirculation as shown in Figure 13. To prevent the diver-

gence of RBC lateral position, experiments were performed to determine a proper number of arrays in regards

to the lateral position of 4 μm diameter beads that model RBCs. Lateral positions of the beads in 4, 6, and 8

arrays were observed at range of Re from 3 to 12, and the contraction region’s length was fixed at 150 μm.

As shown in Figure 14a, it is clear that the lateral position of beads remain near side B at Re = 3 due

to weak drag force. However, as Re increases, the resulting drag force also increases and the beads start to mi-

grate toward side A. The same behavior is observed for the lateral positions in 6 array device until Re = 9 as

shown in Figure 14b. At Re = 12 however, beads migrate back towards side B and results in wider trajectory

band. This occurs because the stream of beads that were already focused near side A are diverged, and they con-

tinue to follow the two counter rotating path of the secondary flow, which recirculates back to side B. Figure

14c shows that the trajectory band is away from side B despite the low Re of 3. This indicates that even though

the magnitude of secondary flow is low, the increased accumulation of displacement due to increased number of

array can cause particles to migrate towards side A. 8 Array device also shows that its trajectory diverges at Re

= 6, which is lower compared to 6 array device, whose divergence occurred at Re = 9.

The average and standard deviation of lateral positions of 4 μm beads in each devices are normalized

by setting the width of the expansion region as 1, spanning from side A to side B. Figure 15 shows the plot of

- 34 -

lateral positions of the beads and it is apparent that trajectory diverges at Re = 9 for 6 arrays and Re = 6 for 8

arrays. This indicates that these devices are not capable of performing high throughput separation compared to 4

array device that can migrate the beads successfully toward side A without diverging. Thus the 4 array is chosen

as a proper array number that can remove RBCs. The average and standard deviation of lateral positions of 4

μm beads in each devices are tabulated by Re in Table 3.

- 35 -

Figure 13. Schematic of particle’s lateral position divergence due to recirculation of secondary flow. Particle’s

lateral positions, after being focused at side A, can diverge after passing through more arrays by entraining in

the secondary flow and recirculating back to side B.

Figure 14. Streamlines formed by 4 μm diameter fluorescent beads captured at the end the of a CEA channel

with a) 4 arrays, b) 6 arrays, and c) 8 arrays.

- 36 -

Figure 15. Plot of the normalized average lateral positions of 4 μm beads in CEA device with 4, 6, and 8 arrays.

Table 3. The average and standard deviations of 4 μm diameter beads in 4, 6, 8 array devices at a range of Re

number from 3 to 12.

Re

Average lateral position

4 Arrays 6 Arrays 8 Arrays

3 0.69 ± 0.16 0.58 ± 0.15 0.43 ± 0.16

6 0.48 ± 0.16 0.30 ± 0.14 0.29 ± 0.19

9 0.36 ± 0.15 0.23 ± 0.14 0.37 ± 0.27

12 0.29 ± 0.13 0.29 ± 0.22 0.46 ± 0.30

- 37 -

4.4. Contraction region length optimization

Unlike the drag force, inertial lift force mainly affects relatively larger particle as mentioned previ-

ously. Inertial lift force in a CEA microchannel can be controlled by changing the height [29] and the length of

the contraction region [30, 32]. In this study the height was held fixed and contraction length was varied to find

a proper condition where white blood cells and cell–bead complex separate from each other. If the length is too

short, they both would migrate to side A by drag force and if the length is too long, they both would migrate to

side B and their trajectories would completely overlap. A proper length would be one that causes cell–bead

complex to migrate to side B as much as possible while U937 cells do not. Based on the measured diameter of

the U937 cells and MCF-7 cells labeled with 5 μm beads, 15 and 25 μm diameter beads were used to predict the

inertial migration of cells in a CEA microchannel. 15 μm is the upper extreme of the U937 cells, which means

that the cell trajectories would be observed between the trajectory formed by 15 μm bead and side A of the mi-

crochannel at the expansion region. 25 μm is the lower extreme of the MCF-7 cells labeled with 5 μm beads,

which means that the cell–bead complex would migrate between the trajectory formed by 25 μm and side B of

the microchannel. Channel height and array numbers were held constant at 55 μm and 4. The lateral positions of

the fluorescent beads were observed at the last expansion region. Figure 16 shows the trajectories of 15 (red)

and 25 μm (green) beads in 150, 300, 450 μm long contraction channels at different Re. In a device with 150

μm long contraction region, 15 μm beads start to migrate toward side A with increasing Re as shown in Figure

16a. However in devices with 300 and 450 μm long contraction region, the lateral position of 15 μm beads stay

near side B, which indicates that inertial lift force dominates the drag force due to the long contraction length.

Figure 16b shows that the trajectories of 25 μm beads stay near side B in all devices for all Re values, indicating

that the large diameter of the bead results in inertial lift force that dominates drag force.

- 38 -

Figure 16. The trajectories of a) 15 μm and b) 25 μm diameter beads observed in CEA devices having 150, 300

and 450 μm long contraction regions. In each devices, the trajectories were observed at a range of Re values

from 3 to 12. Scale bar: 200 μm

- 39 -

The lateral positions of 15 and 25 μm beads in 150, 300, and 450 μm long contraction lengths are

plotted in Figure 17. Whereas 25 μm beads remain at about 0.7 of the expansion channel width with changing

Re, 15 μm bead’s average lateral position approaches side A with increasing Re. The calculated separation reso-

lution, represented by bar graph, also increases with increasing Re value. The separation resolutions calculated

with lateral positions of 15 and 25 μm beads in 300 and 450 μm long contraction region are not as high as those

in 150 μm long contraction region. This indicates that as contraction region’s length increases, 15 μm beads are

exposed to inertial lift force for extended period of time, giving them enough time to migrate back to side B.

From this result, U937 cells are expected to migrate toward side A with increasing Re, while cell–bead complex

will migrate toward side B in a CEA device with 150 μm long contraction region, resulting in size-based separa-

tion.

- 40 -

Figure 17. Plot showing the normalized lateral positions of 15 and 25 μm fluorescent beads in a CEA device

with a) 150 μm, b) 300 μm, and c) 450 μm long contraction region lengths for Re from 3 to 12.

- 41 -

4.5. Optimized channel characterization

From previous results, it is determined that an optimal design for this study consists of 4 arrays with

150 μm long contraction region length. The lateral positions of 4, 15, 25 μm beads that model RBC, WBC and

cell–bead complex respectively, are plotted against Re in Figure 18a. The graph shows that size-based separa-

tion increases with increasing Re, indicating that high throughput sample processing is possible without diverg-

ing of bead lateral position. Figure 18b shows the lateral positions of beads at Re = 12 plotted against the size of

the particles. The Figure shows that the biggest particle migrates to side B due to dominant effect of inertial lift

force and that smallest particle migrates to side A due to dominant effect of drag force.

- 42 -

Figure 18. Graph of a) the normalized lateral positions of 4, 15 and 25 μm beads plotted against Re, and b) the

normalized lateral positions versus particle diameter.

- 43 -

4.6. Characterization of cell trajectories

To observe inertial migration of cells in a CEA mirochannel, different kinds of cells were injected

separately into the device and their lateral positions were imaged at expansion region by using high speed cam-

era as shown in Figure 19. White spheres are identified as U937 and MCF-7 cells (Figure 19a, b). For bead-

labeled MCF-7 cells, small black spots are identified as unlabeled 5 μm beads and round, black spheres are

identified as bead-labeled MCF-7 cells (Figure 19c). The images were analyzed with ImageJ software to meas-

ure the distance from channel wall of side A, and the average lateral positions were normalized.

Figure 20 shows that U937’s average lateral position approaches side A with increasing Re due to

dominant effect of drag force. MCF-7’s average lateral position remains near the center of the channel for all Re,

which indicates that the effect of drag force and inertial lift force are balanced, thus the cells do not migrate to

either side of the channel and instead the standard deviation increases. The average lateral positions values for

U937, MCF-7 cells and bead-labeled MCF-7 cells are tabulated in Table 4. At Re 3 and 6, the two cell’s average

lateral positions are almost the same, which means that separation cannot be performed at this condition. At

higher Re values, the two cells migrate to different lateral positions but their standard deviation overlaps. The

overlap arises from the overlap in U937 and MCF-7 cell’s size distribution and it shows the limitation of size-

based separation of similar sized cells.

However, unlike the unlabeled MCF-7 cells, MCF-7 cells labeled with 5 μm diameter beads result in

average lateral positions farther away from that of U937’s lateral positions at Re = 9 and 12. The distributions of

U937, MCF-7 cells and cell–bead complexes at Re 9 and 12 are shown in Figure 21. The histograms show that

the lateral position distribution of cell–bead complex shifts to side B, or away from that of U937 cells at both

Re values. This indicates that labeling MCF-7 cells with 5 μm diameter beads can reduce the lateral position

- 44 -

overlap between U937 and cell–bead complexes, which would result in better separation.

Furthermore, as shown in Figure 19c, MCF-7 cells labeled with 5 μm beads are observed near side B

and unlabeled 5 μm beads are migrated near side A. Thus, 5 μm beads are expected to flow out through outlet 2,

which means that the unlabeled beads will not be included in the outlet 1 during collection.

- 45 -

Figure 19. Image of a) U937 cells, b) MCF-7 cells and c) MCF-7 cells labeled with 5 μm captured at a wide

section of CEA microchannel after passing through 4 arrays. All images are captured at Re = 12.

Figure 20. The lateral positions of U937, MCF-7 cells versus Re and the lateral positions of MCF-7 cells la-

beled with 5 μm at Re values of 9 and 12.

- 46 -

Table 4. The average and standard deviation values for U937, MCF-7 cells and 5 μm bead labeled MCF-7 cells

at different Re.

Re

U937 MCF-7 5μm bead-labeled MCF-7

AVG SD AVG SD AVG SD

3 0.51 0.016 0.56 0.05

6 0.43 0.08 0.44 0.06

9 0.34 0.10 0.47 0.13 0.59 0.11

12 0.31 0.12 0.46 0.15 0.60 0.14

Figure 21. The distributions of U937, MCF-7 cells and 5 μm bead-labeled MCF-7 cells at a) Re = 9 and b) Re =

12.

- 47 -

4.7. Cell trajectories from mixed sample

Inertial microfluidics based cell separation often experience low separation efficiency at high sample

concentration due to interactions between particles in the channel [23]. The particle–particle interaction can

affect inertial migration of particles, as demonstrated by deterred inertial migration of cancer cells by RBCS in

high hematocrit whole blood [41]. To observe that cell separation can be performed from mixed sample, MCF-7

cell–bead complexes were spiked into a suspension of U937 cells and then their lateral positions were observed

at Re = 12. Because MCF-7 cells were incubated with a solution containing 5 μm diameter streptavidin coated

beads, the spiked sample also contains unbound 5 μm beads. As shown in Figure 22a, white spheres with visible

boundary are identified as U937 cells, small black spots are identified as unbound 5 μm beads, and round black

spheres are identified as cell–bead complexes. The lateral positions of cell–bead complexes and U937 cells are

separated from each other as shown in the Figure 22b. The curve overlap between the lateral positions is smaller

than that of U937 and unlabeled MCF-7 cells. The lateral position of 5 μm beads also does not overlap with that

of cell–bead complex, which means that the unbound beads would be removed from the outlet 2.

- 48 -

Figure 22. a) Image of U937 cells (white sphere), unbound 5 μm streptavidin coated beads (black spots) and

MCF-7 cells labeled with 5 μm diameter beads (black sphere) flowing at Re = 12. b) The distribution of lateral

positions of U937 cells and bead-labeled MCF-7 cells.

- 49 -

4.8. Evaluation of separation efficiency

To separate MCF-7 cells from U937 cells, the wide section of the CEA microchannel has to bifurcate

into two channels that lead to two outlets. Larger cells (MCF-7 cells or bead labeled MCF-7 cells) are to be col-

lected from the outlet 1 (located at side B), and smaller cells (U937) are to be removed through the outlet 2 (lo-

cated at side A). Whether the cells go to outlet 1 or 2 depends on the location of the bifurcation point as shown

in Figure 23. Figure 23a illustrates that placing the bifurcation point near side A causes U937 cells to flow out

through outlet 1, which would result in decreased U937 rejection and decreased MCF-7 purity. On the other

hand, placing the bifurcation point near side B causes MCF-7 cells to flow out through the outlet along with

U937 cells (Figure 23c), which would result in decreased MCF-7 recovery rate. As shown in Figure 23b, it is

important to place the bifurcation point properly in order to collect MCF-7 cells from the outlet 1 and remove as

much U937 cells through outlet 2 as possible.

To determine the optimal bifurcation point, the recovery rate and rejection ratios were calculated for

19 possible bifurcation points with an increment of 0.05 times the width of the channel (0.05 being the closest

to side A and 1.95 being the closest to side B). The cells whose lateral positions are between the wall of side A

and the chosen bifurcation point are considered as removed cells (through outlet 2) and the cells whose lateral

positions are between the bifurcation point and the wall of side B are considered as collected cells (through out-

let 1). The number of cells at each outlet for 19 possible bifurcations were counted and used to calculate the

recovery rate and rejection ratio at each respective bifurcation points.

Recovery rate is the ratio of MCF-7 cells collected from the outlet 1 to total number of MCF-7 cells

that flowed out through both outlet 1 and 2, defined as:

2outlet 1outlet

1outlet

77

7

­MCF­MCF

­MCF

(6)

- 50 -

The U937 rejection ratio is the ratio of U937 cells that are removed through the outlet 2 to the total number of

U937 cells, defined as the following equation:

2outlet 1outlet

1outlet

937937

937

UU

U

(7)

Cells flowing through the wide section of the CEA microchannel were imaged at Re = 12 with high

speed camera as shown in Figure 24. The recovery and rejection ratio curves in Figure 25a, b shows that wheth-

er MCF-7 cells are unlabeled or labeled with beads, placing the bifurcation point too close to side A ensures the

recovery of MCF-7 cells. However, this also allows U937 cells to be collected from the outlet 1, resulting in

low rejection ratio and low purity. On the other hand, placing the bifurcation point too close to side B ensures

complete removal of U937 cells. However, also results in removal of MCF-7 cells and thus the recovery rate

decreases.

The bifurcation point has to be chosen so that both MCF-7 recovery rate and U937 rejection ratio are

the highest. For unlabeled MCF-7 cells, bifurcating channel at 0.35 of the channel width results in the highest

possible recovery rate of 74% and U937 rejection ratio of 76%. However, by labeling of the MCF-7 cells with

beads and bifurcating the channel at 0.5 of the channel width results in increased recovery rate of 97.6% and

rejection ratio of 95%. The altered trajectories of MCF-7 cells caused by bead labeling allows more distin-

guished bifurcation point of separation that results in higher separation efficiency.

- 51 -

Figure 23. Illustration of a) U937 cells not being rejected by placing bifurcation point too close to side A, b)

bead-labeled MCF-7 cells being collected from outlet 1 and U937 cells being removed from the outlet 2 by ide-

al bifurcation placement, and c) bead-labeled MCF-7 cells being removed, resulting in cell loss by placing the

bifurcation too close to side B.

- 52 -

Figure 24. Image of cells flowing through the wide section of the CEA microchannel imaged at Re = 12. a)

U937 cells flowing and reaching the bifurcation near side A (outlet 2). b) Bead-labeled MCF-7 cells flowing and

reaching the bifurcation near side B (outlet 1). Scale bar: 300 μm.

- 53 -

Figure 25. MCF-7 recovery rate and U937 cell rejection ratio curve of a) unlabeled MCF-7 cell separation at Re

= 12 and b) bead-labeled MCF-7 separation at Re = 12.

- 54 -

Chapter 5. Conclusions

5.1 Summary

This study aimed to improve the separation efficiency of a high throughput size-based inertial cell

separation in an inertial microfluidic device by immunospecific size-enhancement of cells with microbeads.

MCF-7 cells were chosen to model circulating tumor cells and were labeled with 5 μm diameter beads to ob-

serve the size-enhancement of MCF-7 cell population. U937 cells were chosen to model WBC population with

large diameter. The sizes of U937 cells and unlabeled MCF-7 cells were measured to compare the size distribu-

tions and their overlap. MCF-7 cells were labeled with 3 and 5 μm beads separately to observe the size en-

hancement. Due to 3 μm bead’s inability to completely overcome the size distribution that overlaps with U937

cells’, 5 μm beads were chosen to enhance the size of MCF-7 cells. Labeling the cells with 5 μm effectively

removed the size overlap between U937 cells and bead labeled MCF-7 cells.

Microfluidic device was fabricated by soft lithography techniques and PDMS microfluidic channel

mold was bonded with glass slide. The device composes of 2 inlets (one for sample and one for PBS buffer) and

2 outlets (one for sample collection and one for waste). Computational fluid dynamics simulation was per-

formed to observe the formation of secondary flow in the microchannel. The simulation results show that in-

creasing the flow velocity results in greater magnitude of the secondary flow. Curl values at the center of sec-

ondary flows were also calculated and the values also increase with increasing flow velocity, which indicates

that the secondary flow’s magnitude increases.

The optimal number of contraction region entrances was determined by observing the lateral posi-

tions of 4 μm diameter fluorescent beads as RBC model. 4, 6 and 8 arrays were tested with fixed contraction

region length and 4 arrays was determined to be the optimal number that would focus RBCs without having

their trajectories diverge.

- 55 -

The optimal length of contraction region was determined by observing the later al positions of 15 and

25 μm diameter fluorescent beads, each of which model U937 cells and bead-labeled MCF-7 cells. The optimal

length of the contraction region was chosen based on the difference in the lateral position of 15 and 25 μm

beads. 150 μm long contraction region was chosen as an optimal contraction length.

In the optimized device, which consists of 4 arrays with 150 μm long contraction region length, U937

cells and MCF-7 cells were injected separately and their lateral positions were observed with varying Re. The

separation resolution between two cells is greatest at the highest Re value tested, which is 12. MCF-7 cells la-

beled with 5 μm diameter beads were then injected into the device and their observed lateral positions were

much farther away from that of U937 cells, which indicates that labeling the MCF-7 cells with beads can effec-

tively improve the separation performance.

The separation performance was judged by comparing the U937 rejection ratio and MCF-7 recovery

of two cases: separation of unlabeled MCF-7 cells from U937 cells; separation of bead-labeled MCF-7 cells

from U937 cells. The calculated recovery rate and rejection ratio were 97.6 and 95% for the separation of bead-

labeled MCF-7 cells, which is higher compared to 74% rejection ratio and 76% recovery rate of unlabeled

MCF-7 separation. Table 5 summarizes the separation performance demonstrated in this study compared to oth-

er inertial size-based separators.

Current sample flow rate reported in this study is 0.88 ml/h, which is based on a single channel. On a

5 inch wafer, maximum of 34 devices can be fabricated and parallelization of 34 devices can theoretically in-

crease the throughput to 29.9 ml/h. The massive parallelization is feasible because it is a continuous separation

device with the ability to collected samples continuously. However for on-chip binding [14] or on-chip captur-

ing [25] methods, parallelization is limited because it results in increased active capture area that needs to be

- 56 -

inspected to locate the captured cells. The process multiplies the time required to locate and analyze the cap-

tured cells by the number of devices that parallelized, which is laborious and increases the work hour of techni-

cians.

Conventionally rare cells have been separated and analyzed by using FACS. FACS however is con-

sidered to be expensive, mechanically complex and require skilled technician to operate [6] as mentioned previ-

ously. The development of high performance microfluidic device has the potential to significantly reduce the

cost of the separation system, miniaturize the work station and eliminate maintenance problems that arise due to

the complexity of the machine. Although the method proposed in this study involves pretreatment of the sample,

considering the fact that separation of extremely rare cells often require magnetic-activated cell sorting (MACS)

to enrich samples [1] and fluorescent labeling, the bead-labeling process can be regarded as a minor problem.

The method can be used to separate and enumerate circulating tumor cells from whole blood for can-

cer screening and used as liquid biopsy to evaluate the efficacy of cancer treatment. The separated cells can also

be collected and their contents can be analyzed for molecular profiling of the captured CTCs. This study in

overall shows promising results of high throughput cell separation with high purity and recovery for many ap-

plications in biomedical research and clinical purposes.

- 57 -

Table 5. Summary of the size-based inertial separation performance demonstrated in recent trends compared to

that achieved in this work.

Hyun, Biosens. Bio-

electron. (2013)

[21]

Sollier, Lab Chip

(2014) [23]

Warkiani, Lab Chip

(2014) [22] This work

Schematics

CTC Re-

covery (%) 93.75 25.8 85 97.6

Purity N/A 100 5 73.8

WBC Re-

jection (%) 90.8 N/A N/A 95

Throughput

(ml/h) 0.6 240 102 0.88

Utilized

forces Inertial lift force Inertial lift force

Inertial lift force

and Dean force

Inertial lift force

and Dean force

- 58 -

5.2 Future directions

Current result shows the separation of bead-labeled MCF-7 cells from U937 cells. It is equivalent to

showing a separation of CTCs from a blood sample after RBC removal. In future study, the MCF-7 cells should

be separated from not only U937 cells, but from a whole blood that contains red blood cells as well, thereby

showing one-step separation of CTC from whole blood.

The MCF-7 cells in this study were labeled with beads separately first, then spiked into a solution

containing U937 cells. Ideally, CTCs should be labeled in whole blood. However, presence of RBCs can signif-

icantly increase the viscosity, which can reduce the chance of beads attaching to CTCs. For labeling in whole

blood to be successful, proper dilution of whole blood may be necessary and possibly more beads should be

added to ensure proper labeling. Once such protocol is developed, further separation studies with blood samples

from cancer patients should also be performed for clinical efficacy.

The size of the beads used in this study was determined by their ability to remove MCF-7 cell’s size

distribution that overlaps with U937 cell’s size distribution. Thus 5 μm beads were chosen over 3 μm beads

simply due to their ability to effectively remove the size-overlap. However, using larger beads and their effects

should also be considered or be tested in future work. Attaching bigger beads with diameters, for example, be-

tween 6 and 10 μm can further increase the size and possibly result in lateral positions that are much farther

away from those of U937 cells. However, in order to use larger beads to cause favorable alteration in lateral

positions, the geometry of the CEA channel, mainly the width of the contraction region, should be modified. If

the width of the contraction region barely allows the bead-labeled cells to pass, it would act as pinched flow

modulation [42] that physically causes cells to order at the center of the channel, which would ultimately result

in lateral positions that are closer to those of U937 cells. Thus, if larger beads are to be used, the width of the

- 59 -

contraction region should be increased accordingly. Furthermore, the size of the beads should not be overly

large in a way that would crush and lyse the cells during incubation process.

Polystyrene beads were attached to cells mainly for the purpose of increasing the inertial lift force

that results from increased diameter. In further studies, magnetic beads can be used instead of polystyrene beads

to not only enhance the effect of size amplification, but also to apply magnetic force for increased separation

efficiency by magnetic selectivity.

Furthermore, beads whose density is higher than that of PBS can be used to observe the effect of

bead-labeled cell’s increased mass. As bead-labeled cells flow through series of curved trajectories, their mo-

mentum can cause them to deflect away from their original streamline [43]. For example, as the bead-labeled

cells enter the contraction region, their momentum would theoretically continue to drive the cells to travel along

the direction of the main flow (along the x-axis) and deflect to the streamline that is closer to side B of the con-

traction region. This may result in increased separation efficiency if the deflection happens such that their re-

sulting lateral positions are further away from those of WBCs.

The method’s current objective is to allow high throughput separation with high purity and recovery

rate for more accurate enumeration and molecular analysis of collected CTCs. Thus this method does not re-

quire cell’s viability, and rather compromise the cell’s viability by bead-labeling. Although it has been proven

from previous literature that bead-labeled cells can be collected and cultured [25], to ensure better attachment of

cells to cell culture dish, use of stripping buffer may be considered. Stripping buffer is typically used in Western

blots to remove antibodies without affecting the proteins that are blotted on paper. Thus, use of stripping buffer

should be tested for bead removal for better cell culturing results. Cultured cells can then be used for screening

of cancer drugs for personalized cancer treatment.

- 60 -

The purity and WBC rejection ratio is mainly affected by few U937 cells that break away from their

original trajectories. Theoretically U937 cells’ trajectories are below the half-width of the expansion region.

However, the trajectories of U937 cells that were observed between the half-width and side B are the ones that

were following behind the bead-labeled cells. Due to the bead-labeled cell’s large diameter, they can disturb the

flow of U937 cells and push them out of their own streamlines. To prevent decreased separation performance

due to particle–particle interaction, the sample can be diluted further or the sample to buffer ratio can be in-

creased in such way that the distance between cells are increased. The decreased throughput caused by diluted

sample can be possibly compensated by integrating parallel channels.

The device’s current throughput is 0.88 ml/h. With further optimization involving increased channel

height and increased cross sectional area of the contraction region, the device is expected to operate with higher

throughput. Few hundreds of microns high channel can be easily fabricated with photolithography and with

proper optimization, much greater amount of sample can be process at a given time.

This study demonstrated the separation of CTCs from WBCs. The principle can also be applied to

separate white blood cells that specifically express surface proteins from other white blood cells. Achieving

such separation by inertial microfluidics based size-separation is impossible because the size difference within

the population of WBC is not significant enough for effective size-based separation to be applied. However, by

enhancing the size of the cells of specific interest with beads, high throughput size-based separation can be

achieved.

Finally, the collected bead-labeled cells can be encapsulated in droplets containing forward primers

for droplet based polymerase chain reaction (PCR) for deoxyribonucleic acid (DNA) analysis. Recent advances

in droplet microfluidics allowed single cell PCR, which provides more accurate analysis results compared to

- 61 -

that of ensemble PCR of a sample that contains multiple cells. The technology encapsulates single cells, func-

tionalized beads that have reverse primers, and dye labelled forward primers [44]. However, the current tech-

nology requires a droplet to contain all three major components for PCR to occur: the cell, functionalized bead

and the primers. The chance having a drop that contains all three components is close to 1 in 100, which means

that the PCR is much less likely to occur for rare cells. By using the method demonstrated in this study, the

beads and cells will always be found together. Thus, by functionalizing the bead not only with antibodies but

also with reverser primers, the chance of PCR to occur will significantly increase. Further studies are required

to integrate the CEA separator with droplet generator; however, when integrated properly, it has the potential to

facilitate single cell droplet based DNA analysis for cancer research.

- 62 -

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Summary in Korean

면역특이적으로 크기가 증대된 혈중 희귀세포의 고순도 회수를 위한

관성 미세유체소자 내에서의 분리법

본 연구에서는 세포와의 마이크로입자 결합을 이용하여 미세유체소자 기반의 혈중암세포

분리기술 효율을 높였다. 혈중암세포를 모델하기 위해 MCF-7 세포주가 사용되었고 백혈구를

모델하기 위해 U937 세포주가 사용되었다. U937 세포주는 백혈구중에서도 크기가 큰 세포들을

모델하기 위해 선택되었으며 각 세포주의 크기 측정을 통해 두 세포주의 크기분포가 겹친다는

사실을 확인하였다. MCF-7 세포표면에 3 혹은 5 μm 지름의 마이크로입자를 부착함으로써

그들의 증가된 크기를 관찰하였고, 5 μm 지름의 마이크로입자 부착을 통한 전체적인

크기증폭은 U937 세포주와의 크기분포오버랩을 제거한다는 사실을 확인하였다.

미세유체채널은 photolithography 기술을 이용하여 제작된 폴리다이메틸실록세인 몰드와

유리 슬라이드로 결합되어 완성되었다. 제작된 미세유체소자는 2 개의 주입구 (샘플 주입구와

버퍼 주입구)와 2 개의 유출구 (샘플 유출구와 폐액 유출구)를 가지고 있으며 주입된 샘플은

확장–축소 영역이 반복되는 미체채널을 통해 흘러가며 크기별로 분리된다. 전산수치

해석방법을 통해 미세채널내에서의 이차유동장 형성을 확인하였으며 유동장으로 인해 발생하는

항력과 이차유동장의 회전속도는 주입되는 유량에 비례하여 증가한다는 것을 확인하였다.

재작된 채널내에서의 형광 마이크로입자의 거동을 통해 세포분리에 최적화된 확장–

축소영의 수와 축소영역의 길이를 결정하였다. 4, 15, 25 μm 지름의 형광 마이크로입자들이

최적화 실험에 사용되었는데 각 크기의 입자들은 적혈구, 백혈구, 그리고 5 μm 입자가 부착된

혈중암세포를 모델하였다.

최적화된 미세유채소자는 4 단의 확장–축소 영역으로 이루어져 있으며 각 축소영역의

길이는 150 μm 로 결정되었다. U937 과 MCF-7 세포주의 거동을 분석한 결과 유속이

높아질수록 미세채널 내에서의 평균 lateral position 차이가 증가하는것을 확인하였다. 5

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μm 입자가 부착된 혈중암세포를 주입하여서 거동을 Re = 9 와 12 에서 분석한 결과 그들과

U937 세포들의 의 평균 lateral positions 차이는 미세입자가 부착되지 않은 MCF-7 세포들과

U937 세포들의 평균 lateral position 차이보다 더 큰것을 확인하였다.

세포 분리 효율은 U937 제거율과 MCF-7 회수율로 평가되었다. U937 세포와 마이크로

입자가 부착되지 않은 MCF-7 의 분리효율은 세포를 따로 주입하여 각각의 lateral positions 을

따로 측정해서 계산되었고 U937 세포와 마이크로입자가 부착된 MCF-7 세포의 분리효율은

같이 혼합된 샘플을 주입하여 관찰된 세포들의 lateral positions 측정을 통해 계산되었다.

마이크로입자가 부착되지 않은 MCF-7 세포의 분리는 74% U937 제거율과 76% MCF-7

회수율을 보여주었지만 마이크로입자를 MCF-7 부착한 경유 U937 제거율은 95%, 그리고

MCF-7 회수율은 97.6%로 증가하였다.

본 연구에서 최적화된 샘플 처리량은 하나의 소자사용을 기준으로 하여 0.88

ml/h 이지만 5 inch 지름의 웨이퍼 하나에서 재작되는 소자의 최대 수인 34 개의 소자를

평행으로 접속시켜 처리량을 29.9 ml/h 로 증가시킬수 있다. 실험에 사용된 분리소자는 세포를

소자 내에 고정시키거나 trap 시키지 않고 연속적으로 분리 및 회수가능여 다수의 소자를

평행으로 접속시킬 수 있다는 장점이 있다. 하지만 칩내에 세포를 고정시켜 분석하는 소자들의

경우 다수의 소자들을 평행 접속시 세포의 위치를 찾는 시간과 분석시간이 그만큼 길어지고

어려워진다는 단점이 있다.

FACS 를 이용한 기존의 세포 분리 방법은 기계가 복잡하고 비용이 많이 들며 숙련된

기술자가 필요하다는 단점이 있다. 하지만 고효율의 미세유체 분리기술의 개발은

세포분리기계의 비용 및 복잡한 기계의 고장으로 인해 생기는 유지비 차감과 연구 작업공간의

소형화 및 쉬운 분리방법을 통해 개선된 세포분리를 가능케 할 수 있다. 본 연구에서 사용된

방법은 분리하고자 하는 세포에 마이크로입자를 부착해야된다는 단점이 있다. 하지만 희귀하게

존재하는 세포분리를 위해 magnetic-activated cell sorting (MACS)를 이용한 샘플 전처리후

FACS 를 이용하는 경우들과 비교해 보았을때 본 연구의 마이크로입자 부착은 작은 단점이라

할 수 있다.

본 연구에서 입증된 분리방법은 혈액내에 있는 혈중 암세포 분리를 통한 암 진단,

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스크리닝 및 생체검사에 응용될 수 있으며, 회수된 암세포는 분자분석 및 유전자 변이

분석에도 이용가능하기 때문에 암 전이 연구뿐만 아니라 세포 분리를 필요로 하는 많은

연구분야에 응용될 수 있을 것이다.

핵심어: 관성 유체역학, 이차유동장, 세포분리, 혈중 희귀세포, 혈중 암세포