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Microfluidic Platforms for Enrichment and Capture of
Circulating Tumor Cells in Blood
Student Researcher: Anthony Han
Faculty Mentor: William C. Tang, Biomedical Engineering Department
There is a plethora of different types of cancer, each with its own distinct causes and
characteristics. However, cancers are generally categorized in two main subgroups:
benign and malignant. The primary difference between the two groups is that malignant
cancers have the ability to spread to other parts of the body during metastasis via blood
vessels while benign cancers do not. If not promptly discovered and removed, most
cancers are deadly due to their potential to metastasize, as cells detach from a primary
tumor site and travel to other parts of the body to form new tumors. These circulating
tumor cells (CTCs) can be detected in the bloodstream.
Metastasis is one of the most dangerous aspects of cancer, as it is often difficult to predict
where the cancer will spread. As such, successful isolation and capture of the CTCs could
potentially aid in the endeavor to promote understanding and knowledge of metastasis.
Furthermore, isolation of CTCs is associated with a wide range of medical and clinical
applications, including monitoring of cancer in patients and medical prognosis. However,
CTC isolation is no simple task as the relative scarcity of CTCs in the bloodstream
compared to other normal hematologic cells makes it difficult to detect and isolate them;
there is approximately 1 CTC for every 1 billion hematologic cells in 1 mL of blood .
There are several approaches possible to achieve CTC isolation. Some of the common
approaches often employ focus on the biochemical properties of CTCs. For example, one
approach is to functionalize the channels of microfluidic devices with antibodies coated
against EpCAM (Epithelial Cell Adhesion Molecule), a protein often overexpressed on
the surfaces of CTCs. However, our research team’s approach is to exploit the
mechanical properties and blood rheology characteristics of CTCs in designing our
device. In essence, we are utilizing shear-modulated inertial microfluidics to separate the
CTCs out from whole blood and to collect them in one of our device’s outlets. This is
possible due to their relatively larger sizes in comparison to normal hematologic cells: as
a comparison, normal cells are ~8-14 μm in diameter while CTCs are ~16-20 μm . We
opt to not use a biochemical approach because we believe that our approach is more
beneficial and advantageous as it allows us to save on the costs of antibodies and reagents
necessary for channel functionalization, thus proving to be more cost-efficient. Another
advantage of our approach is that the CTCs that we collect are viable samples for further
research analysis and studies since they will be preserved in their pristine states, unbound
to and unreacted with any additional reagents nor substrates.
Prior to utilizing inertial microfluidics in our spiral device to isolate CTCs in blood, we
will pass our samples through a preliminary blood debulking stage where we would
essentially filter out the bigger nucleated cells such as CTCs and white blood cells
(WBCs), from the rest of the smaller cellular and protein components of blood. Not only
will this ensure better separation and isolation of CTCs in the later stages, but also further
enriches our sample before it enters the spiral device and increases the purity of the final
The blood debulking method that we will be employing is continuous flow deterministic
lateral displacement (DLD). This stage of the device will be composed of an array of
microposts varying in pillar size and array offset so as to achieve an arrangement of posts
designed to deflect particles above a certain size away from the primary suspension. A
section of the micropost array is shown in Figure 1. In our case, we will be trying to
deflect CTCs and WBCs away from the rest of the blood components, which should not
be deflected by the microposts and would simply pass straight through the array. Thus we
can create two separate outlets for this stage, one designated for the collection of the non-
deflected particles in the primary fluid suspension on the left, and another for the
deflected particles of interest on the right. The deflected particles would then be led to the
inlet of our spiral device where they will undergo the final stages of separation and
collection. As can be seen in the schematic of the array’s functionality in Figure 2, The
WBCs and CTCs will be deflected off to the right while the non-deflected blood
components will pass straight through on the left. 
Figure 1. A section of the micropost array.
Figure 2. A schematic of the
functionality of the micropost
Inertial microfluidics govern the flow profiles for particles in a solution flowing through
curved microchannels. A combination of both an inertial lift force generated by the
channel wall and the Dean Drag Force generated by the Dean vortices arising from
centrifugal acceleration could achieve size-based separation of particles in a solution after
it has reached equilibrium . Thus, CTCs could potentially be separated from other
hematologic cells in blood due to their larger sizes.
The inertial lift force ( ) that a particle of diameter experiences from the channel
walls cause the particle to move away from the walls. It is quantified as follows :
where is the density of the fluid medium, is the fluid shear rate and is proportional
to the maximum fluid velocity within the channel, , such that , with
being the hydraulic diameter of the channel. Also, is the lift coefficient, and is a
function of the position of the particle within the channel regardless of the particle size,
rising from zero at the channel center to a maximum value before falling back to zero
again at a distance of . An illustration of the lift forces acting on a particle
subjected to a parabolic flow through a channel due to the shear gradient and the channel
walls is shown below in Figure 3. 
Figure 3. Illustration of lift forces. Shear induced lift forces are dominant at the
center of the microchannel while wall induced lift forces are dominant near the
The magnitude of the Dean vortices formed due to centrifugal acceleration of the fluid
can be quantified by a dimensionless number known as the Dean number ( ) :
where is the average fluid velocity, is the fluid viscosity, and is the radius of
The particles in a solution experience Dean Drag Forces ( ) that induce the particles to
move laterally as they travel through the curved channels. This force can be expressed as
A graphical representation of the net forces that a particle experiences while flowing
through a curved channel is represented below in Figure 4. 
Figure 4. The net lift forces vertically separate the flowing particles in the
microchannel while the Dean vortices exert a drag force that laterally isolates them.
Both forces contribute to each particle’s unique flow profile and equilibrium
If whole blood were to be run through curved channels such as those in a spiral, the
inertial lift force and Dean drag forces would balance out and cause large cells such as
CTCs to reach an equilibrium position that is close to the proximal end of the curve,
while smaller hematologic cells would equilibrate somewhere near the distal end of the
curve. For this project, we introduce branches on the proximal end of the primary curved
channels to initiate earlier collection of the CTCs. The branches would then divert flow
into another stream that terminates at one of the collection outlets.
Double Spiral Design
Ultimately, our microfluidic device aims to exploit the different flow profiles for particles
of different sizes that are governed by the mechanics of inertial microfluidics, to achieve
size based separation.
The main channel consists of two segments linked in series, each section composing of a
spiral moving in a specific direction (first one counterclockwise and second clockwise).
The main channel is then fixed with 3 collections branches that stem from the inner walls
of the channel and divert flow away from it. The branches on the inner walls are intended
to capture the CTCs/big beads as the flow in these collection branches eventually
congregate and merge into another channel that leads to the CTC collection outlet. The
rest of the solution not collected by the inner branches continues along the primary
channel in the device until they reach the other outlet reserved for hematologic cell
collection. Our device ensures that there will be adequate distance for the solution to run
initially after injection so that the particles in the solution are able to separate and
equilibrate upon reaching the first branch.
Illustrated below, is a mask of our device that we used in fabricating the SU-8 master
mold. Figure 5 illustrates the design for our device which is composed of two spirals each
consisting of 3 whole turns (3 turns counterclockwise followed by 3 turns clockwise).
Figure 5. Mask for our device
The main channel in our device is initially 500 μms in diameter and subsequently
becomes thinner as it passes by the 3 collections branches located on the second spiral
(clockwise) as the solution starts it exit out of the device. Specifically, as the main
channel passes the first branch its diameter is attenuated to 400 μms; 300 μms at the
second branch; 210 μms at the third branch. On the other hand, the collection channel that
the flows diverted away from the main channel by the branches merge into is initially 100
μms in diameter at the first branch, and expands to 200 μms upon the second branch and
290 μms upon the third.
Before we actually move on to test our devices using whole blood, we want to first test
our devices using microbead solutions. The microbead solutions are intended to simulate
whole blood as they are composed of a mixture of fluorescently dyed beads. The
solutions contain green, large beads (~20-27μm) and white, small beads (~7.3μm).
The microbeads solution is supposed to represent whole blood as the large beads are to
represent CTCs while the small beads represent normal hematologic cells. Running this
solution through our device and conducting data analysis on the bead composition of the
outlet streams will allow us to assess the effectiveness and accuracy of our device. After
testing with beads, we hope to move to testing with whole blood, using a spiked number
of cancer cells to test the device.
Our microfluidic devices will be fabricated through soft lithography using
polydimethylsiloxane (PDMS). We will employ CAD softwares such as L-Edit,
Solidworks, and AutoCAD to design the masks that will ultimately be used to pattern the
SU-8 molds of our devices through photolithography techniques; the SU-8 will be
situated on top of silicon wafers. The first step after the molds have been created, is to
hard bake the SU-8 molds so that they retain on the silicon wafers better. Next, we have
the molds undergo a silanization process to prevent the PDMS from sticking to the mold.
After the molds have been preprocessed, we pour the PDMS that we made using a 10:1
ratio of base to curing agent, on top of the molds and allow it to cure at 65°C. We then
extract the device by cutting it out and separating the PDMS from the molds, and then
puncture holes in the device where the inlet and outlets are located. We then clean the
device and prepare it to be plasma treated so that it may irreversibly bond to a glass slide.
The final step is to inspect the devices under an optical microscope to look for defects,
abnormalities, and cleanliness. 
As an undergraduate researcher working on this project in collaboration with my graduate
student mentor, my primary responsibilities involve fabrication of the microfluidic
devices, data acquisition and analysis, and reporting findings to both the research team
and the rest of the lab.
Once I have the device, I can start running the microbead experiment as simulations for
CTC capture. After the bead solution has run through the device, I analyze the results by
using a hemocytometer to count the beads collected at both outlets to assess the accuracy
and success of the device’s ability to capture ‘CTCs’.
Under the guidance of my PI, Professor William Tang, I will be conducting my
experiments in the Microbiomechanics Laboratory located on the third floor of
Engineering Hall. My lab encourages and emphasizes teamwork and I will communicate
my findings and results to the entire lab at the general lab meetings and be open to all
suggestions and feedback.
Review and analyze the results of past research projects and journal
papers to brainstorm and integrate improvements in our new design.
Brainstorm designs for masks for the new double spiral devices on CAD
Conduct experiments on old devices to acquire quantitative data for
future comparison to results of new devices. Establish experiment
protocol. Continue designing masks.
Finish up any experiments and data acquisition for former devices.
Finalize experiment protocol. Finish designing masks for the new
Fabricate new double spiral devices using PDMS and the new SU-8
molds. Run microbeads experiments on the new devices. Assess
accuracy and effectiveness of new designs and if necessary, formulate
design modifications and improvements.
Extract and congregate all data, results, and findings in preparation for
Clean room access fee $650
Fabrication chemicals and supplies $200
Laboratory Consumables $50
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collections from breast cancer patients using immunocytochemical and clonogenic assay
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 Karabacak, N. et al., “Microfluidic, Marker-Free Isolation of Circulating Tumor Cells
from Blood Samples,” Nature, 27, February 2014.
 Zhou, J. Papautsky, I. “Fundamentals of inertial focusing in microchannels,” Lab
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 Chatterjee, A. et al, “Inertial microfluidics for continuous separation of cells and
particles,” SPIE, 7929, 2011.
 Russom, A. et al., “Differential inertial focusing of particles in curved low-aspect-
ratio microchannels,” New Journal of Physics, 11, July 2009.
 Friend, J. R. and Yeo, L., “Fabrication of microfluidic devices using
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