IOP P
UBLISHING
J. Micromech. Microeng.
17 (2007) 1299–1306
J
OURNAL OF
M
ICROMECHANICS AND
M
ICROENGINEERING doi:10.1088/0960-1317/17/7/012
4
1
2
3
3,4
2,4
1 Department of Biomedical Engineering, University of California, Irvine, CA, 92697,
2
USA
Department of Mechanical and Aerospace Engineering, University of California,
3
Los Angeles, CA, 90095, USA
Department of Pediatrics, David Geffen School of Medicine, University of California,
4
Los Angeles, CA, 90095, USA
Institute for Cell Mimetic Space Exploration, University of California, Los Angeles,
CA, 90095, USA
E-mail: lina@uci.edu
, adamyuhlin@gmail.com
, tswong@ucla.edu
, chihming@ucla.edu
, urvashi@ucla.edu
, EMcCabe@mednet.ucla.edu
and jmchen@cmise.ucla.edu
Received 6 March 2007, in final form 14 May 2007
Published 5 June 2007
Online at stacks.iop.org/JMM/17/1299
The ability to generate strong magnetic field gradients is a prerequisite for efficient magnetic-based cell / bio-particle separation or concentration.
Creating these gradients is difficult under microscale fluidic devices.
Conventional MEMS magnetic-based microfluidic devices involve the use of non-trivial and expensive multi-layer fabrication processes in order to produce magnetic field generators / concentrators (e.g. metal coil / ferromagnetic structures) around the microfluidic channels. A microfluidic device with simplified fabrication procedures while achieving the same functional purposes of magnetic separation / concentration of particles is highly desirable. Here, we propose a simple single-layer, single-mask fabrication technique for magnetic MEMS fluidic device construction, where nickel microparticles can be monolithographically integrated into any configurations. We constructed the microfluidic device through conventional PDMS replicate molding, with injection of nickel microparticles into a side channel 25
m apart from the main separation channel. The nickel microparticles are responsible for bending and concentrating the external magnetic field for gradient generation. This magnetic field gradient induced magnetic forces on the particles present in the main channel. The force generated by the presence of the nickel particles is 3.31 times greater than that without the use of a magnetic field concentrator (i.e. nickel particles). The proposed methodology can be extended for the development of automated high-throughput microfluidic cell separation devices. The simplicity of fabrication and enhanced magnetic separation efficiency shows great promise for future microfluidic systems.
(Some figures in this article are in colour only in the electronic version)
1299 0960-1317/07/071299+08$30.00
© 2007 IOP Publishing Ltd Printed in the UK

Y (Adam) Lin et al
Creating large electromagnetic field gradients is crucial for numerous cell or bio-particle separation or concentration applications, such as using dielectrophoresis in achieving
particles separation [ 1 ]. Different from macro-scale devices,
high magnetic field gradients in microscale fluidic system are difficult to generate. Previous developments to generate large magnetic field gradients were achieved by changing the shape and position of magnets that surrounded main fluidic channels. Quadrupole and dipole magnetic systems had been successful to separate cells in channels with diameters
in the millimeter range [ 2 , 3
]. The purity of the separated sample is high (99%) but the recovery rate, defined as the percentage of target cells recovered from the original
sample, is unstable (37–86%) [ 4
].
Recent developments use MEMS technology to generate a magnetic field gradient through the use of micro-coils and magnetic pillars
[ 5 , 6
]. Although these platforms can easily manipulate the magnetic beads in batches, they do not provide a continuous separation.
The above-mentioned MEMS magnetic devices require non-trivial and expensive multi-layers fabrication processes in order to integrate the magnetic materials with the microfluidic channels to achieve magnetic particles separation.
A microfluidic system that allows a simple fabrication procedure while achieving the same functional purpose of magneticbased separation is highly desirable.
Here, we present a simple single-layer, single-mask fabrication technique for magnetic MEMS fluidic devices that is capable to perform cell separation. Generally speaking, magnetic cell separation or manipulation requires a carrier (i.e. magnetic bead) to attach to the target cells. The magnetic beads, also known as
Dynabeads (Invitrogen, CA), are 4.5
µ m superparamagnetic cores with polystyrene shells. The surfaces of the beads are coated with antibodies targeted toward specific cell membrane markers for certain cell types.
Methods for handling the magnetic beads have been very crucial for biochemical
and analytical applications [ 7 , 8 ].
A large interest in cell separation within automated systems has grown among the medical field especially for oncology or hematology research.
Different from the conventional methods of fabricating a magnetic structure within or close to a microfluidic channel to achieve magnetic field focusing, our device exploits a side channel that was located close to the main sample channel
(figure
1 ), where small metal particles can be injected. Small
metal particles, such as nickel, were utilized as the media to concentrate magnetic fields. The presence of the nickel particles in an adjacent side channel increases the magnitude of the magnetic field density gradient which corresponds to an increase in the translational force exerted on the magnetic beads.
Apart from the device fabrication simplicity, this method also has the potential to achieve stable and high recovery rates due to sophisticated force control within the microenvironment.
In addition, the fabrication cost for the device can be relatively low, which may lead to mass production and commercialization for clinical or research purposes.
Figure 1.
A schematic showing the concept of separation of cells / particles attached to the magnetic beads using metal (nickel) particles as a medium to generate a large magnetic field gradient.
The magnetic force, F b
, generated on a magnetic bead is
governed by the following equation [ 9
]:
F b
=
1
2 µ
0
χ · V b
· ∇ B
2
(1) where µ
0 is the magnetic permeability of free space, χ is the difference of susceptibility between the magnetic bead and the surrounding medium, V b is the volume of the bead and
B is the magnetic field density. It is important to recognize that a gradient of the magnetic field density is required for a translational force. A strong uniform magnetic field can only cause rotational force but not translational force.
The total magnetic force acting on a cell with magnetic beads attached is
F m
=
A c
·
α
·
β
·
F b
F d
=
6 π
·
η
· r
· v
(2) where A c is the total surface area of the cell, α is the number of target cell surface markers per membrane surface area, β is the number of antibody bound per marker and F m acting on one magnetic bead.
is the force
Countering the magnetic force is the drag force, F d
, defined by the Stokes drag law:
(3) where η is the viscosity of the medium, r is the radius of the cell and v is the velocity of the cell moving through the medium.
Assuming that gravity and buoyant forces are negligible, the two forces combine into
F m
+ F d
= ma (4) where m is the mass of the cell and a is the acceleration of the cell. The inertial term ( ∼ 10
− 11 ) is several orders of magnitude smaller than the total magnetic force and the Stokes drag force (
∼
10
−
6
) [ 10
]. Thus, we can neglect the inertial term in
equation ( 4 ). This assumption allows us to find the relationship
between the lateral velocity caused by the induced magnetic force and the minimum magnetic field density gradient ( ∇ B 2 ) required.
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Formation of high electromagnetic gradients through a particle-based microfluidic approach
( B )
( C )
( A)
( D )
( E )
Figure 2.
(A) Simulation of the magnetic field density with Ni particles, Ni bar and magnet only. The nickel particles and the nickel bar were placed in between 0 and 50 µ m.
(B) A graph showing the magnetic field density across the center of each simulation case.
(C) A magnified portion of (B) showing the magnetic field density of the center line from 50 to 100 µ m.
(D) The discrete one-dimensional gradient ( B 2 / x ) for each simulation case.
(E) A magnified portion of (D) showing the discrete one-dimensional gradient ( B 2 / x ) between 50 and 100 µ m.
Plugging in equations ( 1 ), ( 2 ) and ( 3 ) into equation ( 4 ),
the relationship between the magnetic field gradient and the lateral velocity of the cell moving in media is obtained:
∇ B
2 =
A c
12 π · µ
0
· η · r
· α · β · χ · V b v.
(5)
By attempting to calculate the relationship between
∇
B
2 and v , the following assumptions were made. First, the number of the magnetic beads bound to each surface marker ( β ) is assumed to be a constant, which, in this case, equals 1 bead / marker.
Second, we assume that the number of markers per area of cell surface ( α ) is also a constant. If one bead is bound to each cell, α equals 8.84
× 10 9 makers m susceptibility of the media ( ∼ 10
− 6
− 2
[ 11 ]. Third, the
) is negligible compared to the susceptibility of the magnetic beads (i.e.
χ
=
1.52)
[ 12 ]. Fourth, the diameter of the cell is between 3
µ m and
10 µ m. We assume that the diameter of the cell is 6 µ m.
Other constants are the permeability of free space, µ
0
10
− 7 N A
− 2 , and the viscosity of media, η = ∼ 10
− 3
= 4 π
×
N s m
− 2 .
The relationships established in this section will be used to estimate the total magnetic force acting on the cell / bead complexes for determining the cell separation efficiency of the device.
To predict the performance of the resulting magnetic separation scheme in the presence of the nickel particles as a magnetic field concentrator, simulations were carried out using a simplified one-dimensional magnetostatic model by commercial software (COMSOL Multiphysics).
Three different scenarios were simulated.
In the first scenario
(case I), a 100 µ m long square magnet with a magnetic strength of 1 T was positioned behind the origin, in the absence of the nickel particles. From the simulation results, the magnetic field decreased dramatically within 100 µ m from the magnet and reached a steady state after 100 µ m from the magnet (figure
2 (B)
). This showed that the maximum force can only be obtained near the magnet (i.e. within 100 µ m from the magnet). To implement this physically, magnets need to be fabricated in extremely close proximity to the sample channel in order for this scheme to be effective for cell separation.
In the second scenario (case II), the nickel particles were put in between the magnet and the fluid to extend the effective range of the magnetic field, and the resulting effects were simulated. The concentration of the nickel particles (of 20 µ m in diameter) used in the simulation was approximately 6
×
10
7 particles ml
− 1
. As shown in figure
2 (A)
, the presence
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Y (Adam) Lin et al of the nickel particles concentrates the magnetic field by bending the field lines. This results in a substantially localized magnetic field gradient, translating in enhanced magnetic
force on the magnetic beads. According to equation ( 1 ), the
force is directly proportional to the gradient of the squared magnetic field density (
∇
B 2 ). Therefore, in order to estimate the amount of force generated on the magnetic beads, the simulated magnetic field data were converted into discrete, one-dimensional gradient values ( B 2 / x ) (figure
2 (D)
).
The ratio between the gradient values with and without the nickel particles indicated that the presence of the nickel particles induces a maximum force (at x
=
50 µ m) that is approximately nine times larger than that in the absence of the particles. At 200 µ m away from the edge of the magnet, this ratio converges to around 3.
Instead of using nickel particles as a magnetic field concentrator, a nickel bar could be used as an alternative to serve the same functional purposes.
Therefore, in the last scenario (case III), a nickel bar, instead of the nickel particles, was placed in between the magnet and the fluid, and the corresponding effects were simulated. The simulation results can be described into two different regimes (i.e. first regime: 0–100 µ m and second regime: > 100 µ m from the origin).
In the first regime, the nickel particles are capable of generating a stronger magnetic field gradient compared to the nickel bar, which translates into higher force generation on the magnetic particles that are being manipulated
(figure
2 (E)
). In the second regime, there are no significant differences between the effects contributed by the nickel particles or the nickel bar (figure
2 (D)
). In addition, one disadvantage of using a nickel bar is that it can only induce a noticeable magnetic field gradient close to its edges, which implies that the effective particle manipulation area is highly restricted.
For example, consider that the adjacent nickel channel is replaced with the nickel bar (figure
1 ), the magnetic
fields would not be bent or concentrated effectively except at the ends of the bar. Therefore, no translational forces can be generated on the magnetic particles passing by the middle portion of the bar. Comparatively, the nickel particles were able to concentrate magnetic fields more efficiently all over the surfaces. Based on the simulation results, we can conclude that using nickel particles as a magnetic field concentrator can be more beneficial in generating enhanced magnetic forces over the nickel bar.
4.1. Channel fabrication
Different channel geometries were designed in conventional computer-aided design software and printed out onto a negative transparency mask (Photoplot, CO). The channels were fabricated using a replicate molding technique. The mold was fabricated using SU-8 negative photoresist (MicroChem,
MA) on a silicon wafer.
The thickness of the mold was
∼
50 µ m. Then, a polydimethylsiloxane mixture (PDMS), in a ratio of curing agent to PDMS at 1 to 10 by weight, was poured onto the mold and subsequently cured at 60
◦
C. After the curing process, the PDMS replicate was peeled off and punched with inlets and outlets at designated locations. To
( A) ( B )
Figure 3.
(A) A mask layout for the microfluidic device. B was the inlet for the sample. A, C and D were the inlets for media. E was the outlet of the waste sample and F was the outlet for the separated sample. G was the inlet for the nickel particles. H was the outlet for the nickel particles. The G–H channel was the adjacent nickel channels for enhanced magnetic field gradient generation.
(B) A schematic illustration showing the corresponding channel dimensions, unit in µ m.
complete the fabrication procedures, both the PDMS channel surface and a glass substrate were activated by oxygen plasma in order to bond the two surfaces together.
All inlets and outlets were 100 µ m in width with the exception of outlet E, which was 150 µ m. The main channel was 200 µ m in width while the adjacent channel was 100 µ m in width. The two channels were 25 µ m apart (figure
3 ). In
addition, a 500 µ l syringe was used at inlet C while 250 µ l syringes were applied for the rest of the inlet locations
(A, B and D) as indicated in figure
4 . Cell and magnetic
beads mixture sample entered the device from inlet B. Cell growth media were inserted from inlets A, C and D. Inlet A was designed to serve the purpose of pushing stagnated cells and beads that were stuck in inlet B into the main channel.
Media from inlets C and D constituted two streams of sheath flows that focused the sample flow into a fine central stream through hydrodynamic focusing. This microfluidic focusing technique allowed us to adjust the position and the width of the sample stream in the same channel design.
4.2. System setup
Following the Dynabead protocol from Invitrogen, 25 µ l of magnetic beads were added to 1 ml of B-lymphocyte sample
(Coriell institute, NJ), at a cell density of approximately
10 6 cells ml
− 1 and mixed for 30 min in a 1.5 ml microcentrifuge tube.
The magnetic beads that are commonly found for analytical purposes were 4.5
µ m in diameter and made
from polystyrene superparamagnetic material [ 13 ].
The
B-lymphocytes were cultured in RPMI 1640 (Mediatech, VA) with 10% FBS and antibiotics 1XPSN (Sigma-Aldrich, MO).
The cells were stained by an addition of 0.5
µ l of Mitotracker red dye (Invitrogen, CA). Roughly 20% volume ratio of glycerol was added to the sample tube to prevent the precipitation of cell / beads complexes in the syringe during
the experiment [ 14 ]. 100
µ l of the prepared mix sample was put in a 250 µ l gas-tight glass syringe (Hamilton, NV) and connected to inlet B. Then growth media were filled into two 250 µ l syringes (connected to inlets A and D) and a
500 µ l syringe (connected to inlet C) (figure
4 ). Once the
setup was completed, the syringes were connected to the microfluidic chip with soft tubing. The chip was placed on
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Formation of high electromagnetic gradients through a particle-based microfluidic approach
Figure 4.
A schematic illustration showing the system connections. The syringes for inlets A and B were placed on one syringe pump
(sample pump) and the other two (C, D) syringes were placed on another syringe pump (media pump). The top small magnet was used in holding the bottom magnet in place.
( A)
( B )
Figure 5.
(A) Locus of the sample stream under the influence of the external magnetic field. The white particles at the bottom of the channel were cells that were pulled out of the stream. This only happened with the presence of the nickel particles. The white dotted lines represent the edges of the main channel.
(B) Locus plot showing the locus of the upper, center and lower bounds of the sample stream. In every ten pixels, the upper and lower bounds of the white stream were taken and averaged. The average of the two created a center line which was line fitted to obtain the first-order coefficient.
an inverted microscope (Nikon TE2000U) that was connected to a CCD camera (AG Heinze, CA). All the fluid media were pumped through digitally controlled syringe pumps (Harvard
Apparatus, MA). The fluid pumping speed for the sample syringe (inlet B), along with one of the 250 µ l media syringes
(inlet A), was set at 0.2
µ l min
− 1 , while the other 250 µ l media syringe (inlet D) and the 500 µ l media syringe (inlet C) were set at 1 µ l min
−
1 .
In order to prove the validity of the increased magnetic field gradient in the presence of nickel particles, three different conditions were tested: (1) in the absence of magnet and nickel particles (termed as cell trial), (2) in the presence of a magnet but without nickel particles (termed as magnet trial) and (3) with the presence of both magnet and nickel particles
(termed as Ni trial). The cell trial was the control experiment that served as a reference to compare with the later results.
Comparison of the magnet trial and the Ni trial determined the contribution of the nickel particles to the magnetic field gradient generation. The magnet in the experiments used was a NdFeB cube magnet with a side length of 4.76 mm (3 / 16 )
(Amazing Magnets, CA). In order to hold the magnet in place on one side of the chip, another small plate magnet, with the dimensions of 3.18 mm × 3.18 mm × 1.59 mm (1 / 8 ×
1 / 8
×
1 / 16 ), was placed on the other side of the chip
(figure
3 ). For the Ni trial, the nickel particles, with less
than 20 µ m in diameter (Atlantic Equipment Engineers, NJ), were immersed in silicone oil that carried the particles into the adjacent side channel from inlet G. The concentration of the particles was approximately 5.31 g ml
− 1 , which corresponds to 1.42
× 10 8 particles ml
− 1 . Fluorescent images were taken at four different locations of the main channel to quantitatively measure the locus of the cell / bead complexes that were subjected to the external magnetic field. At each location,
15 pictures were taken with a 10 s exposure time.
Since the images were taken in four different locations of the main channel, in order to reconstitute the locus of the sample stream, the images were combined using pre-defined alignment points. The images from the first position did not have any usable alignment points; therefore, the images from the other three positions were further analyzed. The pictures from each of the three positions were randomly chosen and linked together to become partial channel images. The images were further processed to enhance the signal-to-noise level for later data analysis purposes (figure
5 ). The fluorescent
images taken were first leveled with an input black and white point of 2 / 1.00
/ 5. The pictures were then joined together and leveled once more to obtain the movement of the cells. Level 2 was set with an input black and white point of 130 / 9.99
/ 132.
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Y (Adam) Lin et al
Figure 6.
A representative set of center locus data from all three trials: Ni trial, magnet trial, cell trial. The starting points of each locus were offset for easier visual comparison.
The locus of the sample stream was traced and drawn from the images. The processed images were used for further data analysis that will be explained in the next section.
The bending of sample stream locus was caused by the force pulling on the magnetic beads attached to the cells. From the center line data of all 15 pictures for the three different trials, the bending of the line from the Ni trial was significantly larger than that of the magnet trial and the cell trial (figure
6 ). In
order to analyze the force exerted on the cell / bead complexes, the velocity values were extracted from the image data to quantify the difference between the three trials. The horizontal velocity of the cell / bead complexes ( V x
) is constant for each experiment since V x depends on the flow rate of the sample and the shear media. On the other hand, the vertical velocity
( V y
) depends on the force exerted on the complexes. From
equation ( 5 ), the total magnetic force is directly proportional
to the lateral velocity of the complexes. Since the vertical
Y range is comparably small, the magnetic force within this range can be assumed to be constant. Therefore, according
to equation ( 5 ), the lateral velocity of the complexes should
be constant. By curve-fitting the locus with equation ( 6 ), the
ratio of the velocity components can be obtained: y
=
V y
V x
· x + y
0
(6) where V x and y
0 and V y are the velocity components of the complexes, is the starting position of the sample stream.
The
first-order coefficient in equation ( 6 ) can be used to quantify
and compare the vertical velocity ( V y
) of the cell complexes in different trials, which can then be used to calculate the magnetic force exerted on the complexes based on equations
( 1
) and ( 5 ).
After running the data through a line fitting function
(MATLAB), the average first-order coefficient over the 15 sets of data for the Ni trial was 8.08
× 10 error of 1.01
× 10 was 2.44
× 10
− 3
− 4
−
3 with a standard while the average for the magnet trial with a standard error of 2.66
× 10
− 4 . The cell trial (i.e. the control experiment) had an average of 1.03
×
10
− 3 with a standard error of 2.57
× 10
− 4 . The percentage errors were 1.2% for the Ni trial, 10.9% for the magnet trial
Figure 7.
The average values for the first-order coefficients for all the three trials.
and 25.0% for the cell trial (figure
7 ). The ratio of the average
first-order coefficients between the Ni trial and magnet trial was 3.31. This directly shows that the presence of the nickel particles has a significant enhancement on the force generated on the cell / bead complexes during the device operation.
In order to conclude the significance of the experimental data, a t -test was performed to verify the statistical difference of our data. The t -value between the Ni trial and magnet trial was 19.79. The t -value between the magnet trial and cell trial was 3.81. The t -value between the Ni trial and magnet trial was
25.55. The p -value for the Ni / magnet trial and the Ni / cell trial was significantly lower than 0.001 for a two-tailed test. For the magnet / cell trial, the p -value was close to 0.01 since a t -value of 2.76 corresponded to a p -value of 0.01 for a two-tailed test.
Overall, the three trials were considered statistically different.
In addition, the effective operational flow rates for the device with the current geometries have been determined by considering the velocity components (i.e.
V x and V y
) of the resulting cell complexes in the sample channel. As mentioned earlier, V x was predominately determined by the shearing flow rate controlled by the media pump whereas V y was determined by the magnetic force induced in a non-homogeneous magnetic field. From the simulation data, the maximum magnetic force acting on the complexes was estimated which was translated into V y
based on equation ( 5 ).
From the current channel geometries, the ratio between V y be at least 5
×
10
−
3 and V x was estimated to in order to have significant lateral
( y -direction) displacement of the cell complexes for effective separation. Therefore, V x and hence the upper bound of the flow rate can be determined indirectly using the simulation data. Based on the calculations, the maximum flow rate for
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Formation of high electromagnetic gradients through a particle-based microfluidic approach the Ni trial is 1.11
µ l min
− 1 and 0.421
µ l min
− 1 for the magnet trial. In the experimental implementation, we have tested the sample flow rate from 0.1
µ l min with 0.1
µ l min
− 1
− 1 to 1 µ l min
− 1 increments and found that 0.2
µ l min
− 1 was a suitable operating flow rate for the current experimental settings. Optimizing the device geometries would increase the maximum flow rate for enhancement of the particle separation throughput while maintaining high separation efficiency.
( A) ( B )
( C) ( D )
The experimental results, in conjunction with the simulation results, have strongly supported our hypothesis that the presence of small nickel particles, close to the sample channel, is able to induce a large magnetic field gradient which translates into an enhanced magnetic force on the cell / bead complexes. The average values for the ratio of the first-order coefficients in the Ni and magnet trials showed that the induced magnetic force in the presence of nickel particles was more than three times stronger than that in the absence of the nickel particles. The results were shown to be statistically different from the t -test. Besides, the percentage errors of the firstorder coefficients of the magnet (11%) and cell trials (25%) showed that the variations among the sample were greater than that of the Ni trial (1%). For the case of the cell trial, the relatively large percentage error was believed to originate from the random diffusion of the complexes or instability of the system such as disturbance from the tubing. In comparison, the small percentage error in the case of Ni trial has shown that a more stable and controllable microenvironment can be maintained in the presence of the nickel particles as magnetic field concentrators.
Despite the successful demonstration of the device for cell separation, one should note that the current device performance is not completely optimized.
A number of improvements can be done to further enhance the device performance. Parameters such as the geometries of the main channel as well as the flow rates for the media and sample are critical in dictating the resulting cell separation performance.
In the experiments, there were two major considerations for determining the current geometrical parameters of the device.
The first consideration was the fabrication issue. In the current design, the gap distance between the sample channel and the adjacent nickel channel is about 25 µ m in width and 50 µ m in depth, which give rise to an aspect ratio of 2:1.
The fabrication of high aspect ratio SU-8 mold requires the use of highly optimized lithography procedures. Therefore, in order to keep the fabrication process simple, we set some design constraints in our devices without compromising too much for the device functionalities. The second reason was about the length scale matching of the cell / bead complexes with the channel geometries. The average diameter of the complex is on the order of 10 µ m. Therefore, in order to prevent blockage of the channel by the cell aggregates, we purposely designed the channel dimensions large enough for the passage of the cell / bead complexes in the medium and at the same time achieving high sample throughput. Although the current geometries may not be the optimized design, it nevertheless was able to demonstrate the working principle of the device for magnetic cell separation successfully. In
Figure 8.
The cell / bead complexes stayed attached to the bottom of the channel and were trapped. The arrows (orange) indicate the flow direction.
(A) The cell / bead complexes were moving along the bottom of the main channel.
(B) The shadow was a group of complexes moving along the bottom of the main channel which completely flowed into the separation outlet channel.
(C) The bottom-left circle was a cell / bead complex moving away from the main stream due to the induced force from the magnetic field gradient generated by the nickel particles.
(D) The cell bead complexes could get stuck at the bottom of the main channel due to the strong magnetic force near the nickel particles.
addition to the channel geometries, the device performance can be adjusted by changing the intensity of the magnetic field gradient through manipulating the nickel particles density on the adjacent channel. Occasionally, some cell / bead complexes may be attracted toward the sidewall of the channel, close to the corner of the adjacent nickel channel (figure
8 ). This can
be prevented by cleverly designing the geometries (e.g. using rounded corners instead of sharp corners) as well as adjusting the location of the adjacent nickel channel.
Practically, our prototyping device has a number of advantages over the conventional micro-magnetic cell separation devices. One of the most significant advantages is the fabrication simplicity which leads to low cost of device production. Recent magnetic bead manipulation platforms require intensive MEMS fabrication technology which are
economically expensive and time consuming [ 5 , 8 , 15 – 17 ].
The fabrication of our device is a simple replicate molding technique which only requires a single-mask layer for the manufacturing process. In addition, the mold can be reused for multiple times to fabricate new channels for testing and optimizing the system.
Optimization is the next step to manufacturing a fully automated magnetic-bead-based cell separation system.
Upon the development of optimized channel designs for maximized cell separation recovery rate and purity, our proposed technique can be extended in producing highthroughput microfluidic cell separation array. For example, an array of the device can be fabricated in a plastic or acrylic cube. These cell separator arrays can be disposable due to their low manufacturing cost. On top of that, the choices of the metal particles are relatively flexible, provided that their permeabilities are large enough for the device to be effective.
The proposed automated system separation device can be further coupled with a microfluidic cell and magnetic bead
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Y (Adam) Lin et al
mixer [ 18 ]. Practitioners using this device are only required
to provide the appropriate magnetic bead and place the sample in the specified container. The separation can then be done automatically. This device will be most useful for researchers who want to study certain cell types or bio-particles from a tissue or blood sample.
Comprehensive Cancer Center gift fund from Wendy and
Ken Ruby.
In this paper, a simple single-layer, single-mask fabrication technique for the construction of magnetic-based MEMS fluidic devices is presented. Compared to the conventional multi-layer MEMS magnetic devices, this new fabrication method offers the flexibility to incorporate small metal particles, such as nickel, monolithically with the fabricated microfluidic device. The particles can bend and / or concentrate magnetic fields to achieve large magnetic field gradients, which in turn produces enhanced magnetic force on magnetic particles. This method can provide a great and economical platform for not only cell separation devices but also many other microfluidic systems.
We would like to thank Dr Branden Brough, Dr Martin Ma and
Nancy Li of the University of California, Los Angeles (UCLA) for the technical assistance. In addition, we would also like to thank Leslie S Liu of the University of California, Irvine (UCI) for the editing of the manuscript. The project was supported by the NIH-funded research project (1R33DK070328) on
‘automated chip-based metabolomic analysis’, the University of California Leadership through Excellence Advance Degrees
Scholarship Program (UCLEADS) and UCLA’s Jonsson
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