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Two-stage microfluidic chip for selective isolation of circulating tumor cells (CTCs) Kyung-A Hyun a,1, Tae Yoon Lee b,1, Su Hyun Lee c, Hyo-Il Jung a,c,n a
School of Mechanical Engineering, Yonsei University, 262 Seongsan-no, Seodaemun-gu, Seoul 120-749, Republic of Korea Materials & Components R&D Laboratory, LG Advanced Research Institute, Seoul 137-724, Republic of Korea c Nano Medical National Core Research Center, Yonsei University, 262 Seongsan-no, Seodaemun-gu, Seoul 120-749, Republic of Korea b
art ic l e i nf o
a b s t r a c t
Article history: Received 30 May 2014 Received in revised form 9 July 2014 Accepted 9 July 2014
Over the past few decades, circulating tumor cells (CTCs) have been studied as a means of overcoming cancer. However, the rarity and heterogeneity of CTCs have been the most significant hurdles in CTC research. Many techniques for CTC isolation have been developed and can be classified into positive enrichment (i.e., specifically isolating target cells using cell size, surface protein expression, and so on) and negative enrichment (i.e., specifically eluting non-target cells). Positive enrichment methods lead to high purity, but could be biased by their selection criteria, while the negative enrichment methods have relatively low purity, but can isolate heterogeneous CTCs. To compensate for the known disadvantages of the positive and negative enrichments, in this study we introduced a two-stage microfluidic chip. The first stage involves a microfluidic magnetic activated cell sorting (μ-MACS) chip to elute white blood cells (WBCs). The second stage involves a geometrically activated surface interaction (GASI) chip for the selective isolation of CTCs. We observed up to 763-fold enrichment in cancer cells spiked into 5 mL of blood sample using the μ-MACS chip at 400 μL/min flow rate. Cancer cells were successfully separated with separation efficiencies ranging from 10.19% to 22.91% based on their EpCAM or HER2 surface protein expression using the GASI chip at a 100 μL/min flow rate. Our two-stage microfluidic chips not only isolated CTCs from blood cells, but also classified heterogeneous CTCs based on their characteristics. Therefore, our chips can contribute to research on CTC heterogeneity of CTCs, and, by extension, personalized cancer treatment. & 2014 Elsevier B.V. All rights reserved.
Keywords: Circulating tumor cell (CTC) Selective isolation Microfluidic magnetic activated cell sorting (micro-MACS) Geometrically activated surface interaction (GASI) Epithelial cell adhesion molecule (EpCAM) Human epidermal growth factor receptor 2 (HER2)
1. Introduction Tumor heterogeneity has been observed between tumors of the same histopathological subtype (inter-tumor heterogeneity) or within a primary tumor and its metastases and even within a single tumor at either a primary or a metastatic tumor site (intratumor heterogeneity) (Hayes and Paoletti, 2013; Fisher et al., 2013). Inter-tumor heterogeneity is well established in breast and colon cancers and other solid tumors (Cusnir and Cavalcante, 2012). According to the characteristics of tumor subtypes, remarkable differences exist in prognosis and therapeutic strategies (Goldhirsch et al., 2011). In the latter case, intra-tumor heterogeneity may be important in personalized medicine approaches where single tumor biopsy samples are generally needed to portray tumor mutational landscapes (Gerlinger et al., 2012). n Corresponding author at: School of Mechanical Engineering, Yonsei University, 262 Seongsan-no, Seodaemun-gu, Seoul 120-749, Republic of Korea. Tel.: þ 82 2 2123 5814; fax: þ82 2 312 2159. E-mail address: [email protected] (H.-I. Jung). 1 These authors contributed equally to this paper.
In this light, circulating tumor cells (CTCs) have been considered a viable and readily accessible alternative source of tumor cells in the form of a “liquid biopsy” to clarify the intra-tumor heterogeneity (Cristofanilli et al., 2004). Nevertheless, CTC research has been challenged by the rarity (1–100 CTCs per 1 mL of peripheral blood from a human cancer patient) and heterogeneity of CTCs. Numerous approaches with microfluidics have been developed to obtain pure CTCs from a large number of normal hematological cells. The two main approaches are positive and negative enrichment. The positive enrichment method specifically isolates the target cells (CTCs) from non-target cells (normal hematological cells) using their specific characteristics, such as size, electric charge and surface protein expression, while the negative enrichment method specifically elutes non-target cells (Hyun and Jung, 2014). Each method has its own advantages and disadvantages. The positive enrichment can separate CTCs with high purity, but can be biased by the selection criteria and may result in a loss of heterogeneous CTCs which fail to meet the desired criteria (Hughes and Mattison, 2012; Kim et al., 2013; Ozkumur et al.,
http://dx.doi.org/10.1016/j.bios.2014.07.019 0956-5663/& 2014 Elsevier B.V. All rights reserved.
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2013). Although negative enrichment can isolate intact and heterogeneous CTCs, the purity is generally lower than that of positive enrichment (Hyun et al., 2013a, 2013b; Ozkumur et al., 2013; Chen et al., 2011). Additionally, because of their rarity, CTCs must be enriched from a large volume of blood (typically 7.5 mL) to obtain reliable information. Microfluidic devices can continuously isolate CTCs with lower cell loss than a batch process and manipulate small amounts (10 6–10 12 L) of fluid. Thus, along with recovery of heterogeneous CTCs and purity, throughput of microfluidic devices is also an important factor for assessing the performance of CTC isolation. Although several microfluidic devices can isolate various CTCs and then verify their heterogeneity such as the expression ratio of epithelial and mesenchymal cell markers, the research only identifies the existence of heterogeneous CTCs without reseparating the cells based on their characteristics (Hyun et al., 2013b; Yu et al., 2013). Herein, we present a two-stage microfluidic chip for selective isolation of CTCs. The two-stages consisted of two types of microfluidic chips; 1) a microfluidic magnetic-activated cell sorter (μ-MACS) which uses planner microfluidic channels and magnetic fields to deplete magnetically labeled cells from the mixed cells in a continuous flow and, consequently, achieves a highly efficient sorting of the target cells and 2) a geometrically activated surface interaction (GASI) chip which exploits enhanced interactions of target cells and the surface inside the microfluidic channel. In the first stage, the magnetic nanoparticle-coated white blood cells (WBCs) were eliminated in the magnetic field whereas nonlabeled CTCs were freely released into the next stage. In the second stage the released CTCs were captured in the GASI chip depending on their expression level of epitope (e.g. EpCAM or Her2). Our experimental results showed the two-stage microfluidic chip is very useful for selectively and specifically isolating CTCs having varying surface protein expression.
2.2. Immobilization of captured molecules in the GASI channel EpCAM or HER2 antibodies were used to capture EpCAM positive cells, or HER2-positive cells inside the channel, respectively. To coat the channel with these antibodies, an affinity surface was prepared using established protocols with slight modifications. First, a solution of 0.2 mg/mL neutravidin in 10 mM Tris buffer was introduced into the channel using a syringe pump (flow rate: 100 μL/min), followed by incubation for 1 min and then rinsing with 10 mM Tris buffer. Then, each biotinylated antibody was added to complete the surface conjugation by forming the avidin–biotin complex and incubated for 10 min, followed by two rounds of rinsing with phosphate buffered saline (PBS) containing 1% bovine serum albumin (BSA) to prevent nonspecific binding. All reagent incubations were performed at room temperature. 2.3. Sample preparation Various human breast cancer cell lines (MCF-7, SK-BR-3, MDAMB-231) were used as the CTC model to evaluate the performance of the chips. The cells were cultured at 37 °C with 5% CO2. After detachment from the plate by trypsin treatment, the cells were resuspended in 10 mL PBS containing 2% fetal bovine serum (FBS). Human blood samples (5 mL) from a healthy volunteer were collected in vacutainer tubes containing EDTA and processed within 6 h. For the elimination of erythrocytes, we used hemolysis buffer (Qiagen, Chatsworth, CA, USA), which was added to the blood samples in a 1:10 v/v ratio and incubated for 10 min at room temperature. The samples were centrifuged at 600g for 10 min and the pellets resuspended in 3 mL of PBS containing 2% FBS. EasySeps human whole blood CD45 depletion cocktail (StemCell Technologies, North America, www.stemcell.com) was added to the samples in a 1:10 blood volume/cocktail volume ratio and incubated for 15 min to coat the CD45 antibody-conjugated magnetic nanoparticles to leukocytes. EasySeps magnetic nanoparticles were added in sequence as described above and the mixture was incubated for 10 min.
2. Experimental section 2.4. Cell staining 2.1. Microfluidic chip fabrication A microfluidic chamber was fabricated using a laser machine (Universal Laser Systems, Scottsdale, AZ, USA) for rapid prototyping. The laser machine cut the designed microfluidic chamber from a PMMA sheet that was sandwiched between double-sided tapes (3M 9475LE; 3M Company, St. Paul, MN, USA). The top and bottom protective layers of the cut PMMA sheet were peeled away revealing the permanent adhesive surfaces. Then, each laser-cut polyester film (3M Hydrophilic Polyester Film 9962; 3M Company, St. Paul, MN, USA) was pressed on the top and bottom of the permanent adhesive surface. The chamber height was 930 mm and total volume 1 mL. To create a magnetic field gradient for efficient attraction of magnetic nanoparticles, NdFeB block magnets (EP331, EP332; e-Magnets UK, Hertfordshire, UK) with a maximum energy product of 45 MGOe (grade N45) were arrayed in a laser-cut plastic cartridge (Fig. S1). A strong magnetic field was generated between the magnetic field lines of two adjacent magnets. The arrangement was parallel to the flow direction and provides magnetic nanoparticles with longer exposure time within the induced magnetic field gradient. The microfluidic chamber was sandwiched between the magnet array cartridges to increase the possibility of capturing leukocytes labeled with magnetic nanoparticles. Conventional two layer soft lithography was used to pattern the focusing þGASI channel (Hyun et al., 2013b).
The cells collected from the outlet were fixed on a glass slide using cytospin and subsequently stained with immunofluorescence for DAPI, cytokeratin, EpCAM and CD45 to identify nuclei, CTCs, epithelial cells and leukocytes, respectively. Finally, the slide glass was fitted on an automated image analysis system and photographed by a moving stage linked with an image-taking system programmed by MATLABs. The cells were labeled with the PKH26 red fluorescent cell linker kit purchased from Sigma-Aldrich according to the manufacturer's protocol with minor modifications. Briefly, the cell suspension was washed with PBS twice (centrifugation was performed at 400g for 5 min) and the cells were suspended in 1 mL of Diluent C (Sigma-Aldrich). Simultaneously, 4 μL of PKH26 dye (Sigma-Aldrich) was added to 1 mL Diluent C and incubated with the cell solution for 5 min at room temperature. To remove the excess dye, 2 mL of 1% BSA in PBS was added and incubated for 1 min. Cells were then resuspended in 10 mL of complete media and washed three times. 2.5. Operation We used a syringe pump (KDS210, KD Scientific Inc., MA, USA) to inject the prepared sample continuously into the channel. The flow rate in the μ-MACS chip was from 100 to 400 μL/min and in the GASI chip from 80 to 100 μL/min. Prior to μ-MACS chip
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experiment, we performed a degassing process by filling the channel with PBS with 2% FBS to prevent non-specific binding of cells. To observe the enrichment efficiency after the process, the cells collected from the outlet were spin-coated on a slide glass at 1500 rpm for 5 min. An image of the entire area was captured using auto-stage coupled fluorescence microscope. Using the target cell identification criteria, the captured images were examined.
3. Results and discussion 3.1. Design and function of the two-stage microfluidic chip Previously, we introduced a GASI chip to isolate intact and heterogeneous CTCs (Hyun et al., 2013b). This technique employs an asymmetric herringbone that generates enhanced mixing flows, increasing the surface interaction between the cells and antibody-conjugated channel surface. However, due to low operational flow rate (20 uL/min), this chip was not applicable for the process of fully drawing blood in a clinical setting, typically 7.5 mL. Before injecting the blood samples into the GASI chip, the cells were preconcentrated into 1–2 mL of buffer by conventional centrifugation to handle the large sample volume, but the capture efficiency with a high concentration of cells decreases because the captured cells tend to detach by interaction with cells traveling in the flow stream (Sin et al., 2005). Additionally, although this chip can isolate heterogeneous
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CTCs, the cells need to be sorted based on their characteristics to investigate their heterogeneity. Therefore, we developed a new two-stage enrichment chip to improve the throughput and selectively isolate the CTCs based on their surface marker. In the first stage (μ-MACS), WBCs coated with CD45 antibody-conjugated magnetic nanoparticles were captured inside the chip by magnetic force and heterogeneous CTCs exited through the outlet. After the first-stage of enrichment, all the cells were congregated in the center of the channel by inertial force and then moved into the GASI region. When the cells entered the GASI region, the EpCAM-positive cells were specifically captured on the EpCAM antibody-coated channels, while the EpCAM-negative cells were eluted through the outlet (Fig. 1). Currently, the conventional batch MACS system is too bulky in size to carry for point-of-care medical systems (Chen et al., 2014) and cannot provide a continuous process necessary to eliminate the risk of losing target cells (Hyun and Jung, 2014). Therefore, microfluidic technology is needed to provide portability and continuous flow for more precise control of the magnetic and hydrodynamic forces acting on the target cells to improve separation. The μ-MACS chip consists of an inlet, an outlet, four seriallyconnected straight channels, and two neodymium magnetic plates. Because the magnetic flux decreases dramatically with distance from the magnet surface (Pamme, 2006), thin polyester films (1-mm-thick) were used as fluidic channel covers, which bordered both the top and the underside of the magnet to increase the bias of magnetic nanoparticle-tagged cells to the magnetic force. The magnet plate was composed of 21 rectangular
Fig. 1. Schematic diagram of a two-stage enrichment chip for selective isolation of CTCs. The two-stage microfluidic chips consisted of a microfluidic magnetic activated cell sorter (μ-MACS) and an inertial focusing channel combined with a geometrically activated surface interaction (GASI) chip. After the μ-MACS chip eliminated a number of white blood cells (WBCs) via negative enrichment, the GASI chip was used for positive enrichment to selectively isolate cells based on their surface protein expression.
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neodymium magnets. The two neodymium magnetic plates were assembled to the microfluidic channel by sandwiching. Although 99% of WBCs were eliminated with the first-stage μMACS chip, a large sample volume remains a throughput issue for the GASI chip. Therefore, we devised a strategy to enhance the throughput of the GASI chip by adding a focusing channel. Prior to entering the GASI region, the sample flow was decelerated by the inertial focusing channel. In a rectangular microchannel, the shearinduced lift force drives particles toward the channel walls while the wall-induced lift force drives particles toward the channel centerline. The microparticles migrate to the equilibrium position by the balance of the two lift forces. In particular, particles focus into two equilibrium positions centered at the top and bottom walls at a low aspect ratio (AR ¼height/width) microchannel at a moderate Reynolds number (Re), which is expressed by the following equation:
Re =
ρUmDh μ
where ρ is the density of the fluid, Um is the maximum flow velocity in the channel, Dh is the hydraulic diameter of the channel, and μ is the dynamic viscosity of the fluid. Reynolds number is a dimension less number that prioritizes between inertial force and viscous force in the microchannel (Zhou and Papautsky, 2013). Thus, the focusing channel was designed to draw all the cells into the GASI region and discard a portion of the buffer to the outlets as waste consequently, the GASI region adjusted by changing channel dimensions. The flow rate was directly proportional to the pressure drop in the microchannel because it was the product of the cross-sectional area of the channel and average velocity. As mentioned above, because the GASI operated at a low flow rate, we determined the channel dimensions for the ratio of waste outlet flow rate to the GASI region to be nearly 7:3. The height and width of focusing channels were 40 and 100 μm, respectively (AR ¼0.4) and the focusing channel length was 14 cm. The width of waste channels was 150 μm, while the width of channel toward the GASI region was 200 μm. 3.2. Determination of experimental conditions and evaluation of chip performance We determined the experimental conditions for the first-stage negative enrichment process (μ-MACS chip). To investigate the depletion limit of the μ-MACS chip, we prepared three different volumes of hemolyzed blood samples from healthy donors (3, 5,
and 7.5 mL) resuspended in 3 mL of PBS with 2% FBS and then conjugated with magnetic nanoparticles. We loaded the samples into the inlet and then counted the number of cells that exited the outlet using a hemocytometer. The depletion efficiency was over 99% regardless of cell concentration and flow rate because the magnetic flux density of the neodymium magnets was sufficient for capturing the magnetic nanoparticle-coated leukocytes (almost 1700 G per neodymium magnet) (Fig. 2a). In CTC research, recovering as many target CTCs as possible is also important due to their rarity. When the leukocytes and cancer cells were loaded at a fixed concentration of 1 × 106/mL into the inlet, 90.97% of cells were recovered at 400 μL/min inlet flow rate (Fig. 2b). Although the high flow rate dictates the high throughput performance of a chip, sample leakage was observed at the upper flow rate (400 μL/min) because of the high pressure. Thus, we determined the optimal flow rate as 400 μL/min. The number of cancer cells (MCF-7) isolated using the μ-MACS chip was proportional to the number of cancer cells spiked into blood samples across the physiologically relevant range (50–1000 cancer cells/5 mL of human blood) (Fig. 2c). The enrichment yield was as high as 763.14 fold. In the second stage (GASI chip), the enriched cells on the μMACS chip were reseparated based on their surface markers by capture with specific antibody-coated GASI channels. Regarding a recent study reporting that CTCs and WBCs overlap in size (Kim et al., 2012), we considered all the cells as CTCs because the purpose of the second stage enrichment was reclassification of cells, not elimination of leukocytes. Thus, the green 7- and red 15μm fluorescent beads were tested at a fixed concentration of 1 105/mL to ensure all the cells congregated into the center of the channel. Although the focusing efficiency increased as the Re increased in the low aspect ratio microchannel, increasing the flow rate indefinitely was difficult because the high shear force militated against affinity-based cell capture in the GASI region. Therefore, we designed a sufficiently long focusing channel to concentrate the small microbeads into the channel center at a moderate Re (Re¼38–47). Within the focusing channel, relatively larger 15μm beads were well focused across all 80–100 μL/min flow rates (from 97.3571.98% to 98.23 73.54%; n ¼3). Regarding the 7-μm beads, the focusing efficiency increased from 87.18 74.17% to 95.01 71.95% as the flow rate increased (Fig. 3a). Therefore, we determined the optimal flow rate to be 100 μL/min. Fig. 3b shows the behavior of fluorescent microbeads in the focusing channel and outlets at the 100 μL/min flow rate. When three different concentrations of hemolyzed blood samples and cancer cell
Fig. 2. Determination of optimal conditions. (a) Depletion efficiency depending on the flow rate and concentration of blood cells. The depletion efficiency was almost 100% with these terms. (b) Effects of flow rate on recovery. The recovery was under 10% at a flow rate from 100 μL/min to 200 μL/min and increased to approximately 90% at a flow rate of 300 μL/min. (c) Regression analysis of capture yield for a range of MCF-7 cell concentrations spiked into whole blood. The number of separated cells was proportional to the increased number of spiked cells (y = 0.913x + 8.3888, R2 = 0.9991) .
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Fig. 3. (a) Focusing efficiency based on flow rate and bead size. Flow rates from 80 to 100 mL/min did not cause any significant differences in efficiency. However, the focusing efficiency with the 15-μm bead was higher than with the 7-μm bead. (b) Fluorescent images illustrating migration of 7-μm and 15-μm beads toward the microchannel center. Flow is from left to right. (c) Focusing efficiency based on the number of blood cells.
samples were injected individually into the inlet, the 100-fold diluted blood sample reached almost 100% focusing efficiency and 95.39% of cancer cells were successfully focused (Fig. 3c). While the focusing efficiency was relatively lower with undiluted than diluted blood samples, this was an acceptable value because a large amount of blood cells were eliminated by first-stage negative enrichment. 3.3. Selective separation of cancer cells To verify that the GASI chip selectively separated the CTCs based on their surface markers, we evaluated the separation yield using three different breast cancer cell lines (MCF-7, SK-BR-3 and MDA-MB-231 cells). EpCAM surface densities on the cell lines ranged from 2.22 105 binding sites/cell on MCF-7 to 1.7 103 binding sites/cell on MDA-MB-231. The HER2 surface densities ranged from 9.76 105 binding sites/cell on SK-BR-3 to 14.7 103 binding sites/cell on MDA-MB-231 (Prang et al., 2005). The results indicated the MCF-7, SK-BR-3 and MDA-MB-231 cells can be considered as EpCAM-positive, HER2-positive, and both EpCAMand HER2-negative, respectively. We successfully separated tumor cells based on their EpCAM and HER2 expression levels on the anti-EpCAM antibody-coated chip and anti-HER2 antibody-coated chip, respectively (98.81% of MCF-7 cells (EpCAMþ ) and 93.12% of MDA-MB-231 cells (EpCAM ), 86.51% of SK-BR-3 cells (HER2 þ) and 66.54% of MDA-MB-231 (HER2 )) (Fig. 4). It was thought that the reason why the EpCAM based separation yield is higher than that of HER2 based separation is that the difference of EpCAM expression level between MCF-7 and MDA-MB-231 cells (approximately 130 times) is greater than the difference of HER2
expression level between SK-BR-3 and MDA-MB-231 cells (roughly 66 times). In clinical samples, the heterogeneity of cells and their fractions remain undetermined; therefore, mixed cell experiments were conducted to show the GASI chip can handle any fraction of samples. When the mixed MCF-7 cells and MDA-MB-231 cells were injected into the anti-EpCAM antibody-coated GASI chip, the cells were sorted according to their surface markers, even if the smaller ratio of MCF-7 and MDA-MB-231 cells induced a lower purity in MCF-7 cells similar to the inverse ratio. The mixed cell experiments with SK-BR-3 and MDA-MB-231 cell lines produced similar results (Table 1). Finally, we tested the two-stage enrichment by separating cancer cell lines spiked into 5 mL of blood samples. Sample 1 was prepared by spiking 103 of MCF-7 cells (green cell tracker staining) and 103 of MDA-MB-231 cells (red cell tracker staining) per 1 mL of blood and sample 2 was prepared from 103 of SK-BR-3 cells (green cell tracker staining) and 103 of MDA-MB-231 cells (red cell tracker staining) at the same concentration as sample 1. The WBCs in these two samples were eliminated by the μ-MACS chip and then sample 1 was injected into the EpCAM antibodycoated GASI chip while sample 2 was infused into the HER2 antibody-coated GASI chip. After the samples passed through the GASI chip, the separated cells were fixed on a slide glass and stained with DAPI for DNA content. After imaging the slides, the captured images were examined according to predefined criteria (MCF-7 and SK-BR-3 cells: DAPI positive, green fluorescent positive; MDA-MB-231 cells: DAPI positive, red fluorescent positive; WBCs: DAPI positive, green and red fluorescent negative) (Fig. 5). Because the GASI region is too large to take pictures, while the
Fig. 4. Second-stage (GASI chip) enrichment for selective capture of CTCs. Tumor cells were selectively separated based on their EpCAM and HER2 expression levels on a GASI chip (98.81% of MCF-7 cells (EpCAM þ ) and 93.12% of MDA-MB-231 cells (EpCAM ), 86.51% of SK-BR-3 cells (HER2 þ) and 66.54% of MDA-MB-231 cells (HER2 )).
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Table 1 Comparison of purity according to the ratio of mixed cancer cells. Ratio (MCF-7/MDA-MB231) [SK-BR-3/MDA-MB231]
MCF7 Purity (%) [SK-BR-3 Purity]
MDAMB231 Purity (%) [MDA-MB-231 Purity]
1:1 1:10 1:100 10:1 100:1
50.63 [55.67] 9.61 [9.80] 0.24 [1.04] 91.10 [91.09] 99.43 [99.14]
80.39 [71.98] 96.85 [94.66] 91.10 [99.24] 11.04 [10.35] 5.22 [1.90]
cells passed through the outlet is spin-coated in 1 mm diameter circle area, the MCF-7 and SK-BR-3 cells (yellow dashed circles) appeared fewer than MDA-MB-231 cells (white dashed circles). The purities used for measuring the purity of the target cells were 22.91% for MCF-7 and 15.43% for MDA-MB-231 in EpCAM-based selection, 13.37% for SK-BR-3 and 10.19% for MDA-MB-231 cells in HER2-based selection. Although a small amount of WBCs remained, the purity of the spiked cancer cells increased significantly to 22.91% in the two-stage microfluidic chip (μ-MACS chip and GASI chip) compared with 3.64% in the GASI chip.
Additionally, our two-stage microfluidic chip can be used as two-stage negative enrichment to produce high enrichment yields. When the cells were injected into the first stage μ-MACS chip and then second stage GASI chip, the negative enrichment ratio increased by approximately 44 106 fold (data not shown).
4. Conclusion CTCs have been certified as diagnostic and prognostic indicators of cancer in many clinical studies with the help of the many techniques that have been introduced for CTC isolation. However, the heterogeneity of CTCs, which may be important in personalized cancer medicine and for the development of new anti-cancer drugs has not been determined. In this study, we constructed a high-throughput and high-purity two-stage microfluidic chip for WBC elimination and selective CTC isolation (Table S1). The first stage (μ-MACS) was a high throughput negative enrichment for WBCs elimination via magnetophoresis. In the second stage (specific antibody-coated GASI chip), the cells collected from the μ-MACS chip were selectively separated based on the affinity of specific antibodies for their surface proteins. Because antibodies
Fig. 5. Fluorescent images of cells selectively isolated using a two-stage microfluidic chip. Each cell was identified by predefined criteria (MCF-7 and SK-BR-3 cells: DAPI positive, green fluorescent positive; MDA-MB-231 cells: DAPI positive, red fluorescent positive; WBCs: DAPI positive, green and red fluorescent negative) (For interpretation of the references to color in this figure, the reader is referred to the web version of this article).
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can be selected based on the research purpose, our two-stage microfluidic chip was designed to provide various opportunities for cell heterogeneity investigation in rare cell research including CTCs. When the 3-mL sample was used, the entire process from the μ-MACS chip to the GASI chip was completed in 1 h.
Acknowledgments This study was supported in part by a grant from the Korea Health Technology R&D Project (Grant no. A121986) of the Ministry of Health & Welfare and the Center for BioNano Health-Guard funded by the Ministry of Science, ICT & Future Planning (MSIP) of Korea as Global Frontier Project (H-GUARD_2013M3A6B2078959).
Appendix A. Supplementary information Supplementary data associated with this article can be found in the online version at http://dx.doi.org/10.1016/j.bios.2014.07.019.
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Biosensors and Bioelectronics journal homepage: www.elsevier.com/locate/bios
Two-stage microfluidic chip for selective isolation of circulating tumor cells (CTCs) Kyung-A Hyun a,1, Tae Yoon Lee b,1, Su Hyun Lee c, Hyo-Il Jung a,c,n a
School of Mechanical Engineering, Yonsei University, 262 Seongsan-no, Seodaemun-gu, Seoul 120-749, Republic of Korea Materials & Components R&D Laboratory, LG Advanced Research Institute, Seoul 137-724, Republic of Korea c Nano Medical National Core Research Center, Yonsei University, 262 Seongsan-no, Seodaemun-gu, Seoul 120-749, Republic of Korea b
art ic l e i nf o
a b s t r a c t
Article history: Received 30 May 2014 Received in revised form 9 July 2014 Accepted 9 July 2014
Over the past few decades, circulating tumor cells (CTCs) have been studied as a means of overcoming cancer. However, the rarity and heterogeneity of CTCs have been the most significant hurdles in CTC research. Many techniques for CTC isolation have been developed and can be classified into positive enrichment (i.e., specifically isolating target cells using cell size, surface protein expression, and so on) and negative enrichment (i.e., specifically eluting non-target cells). Positive enrichment methods lead to high purity, but could be biased by their selection criteria, while the negative enrichment methods have relatively low purity, but can isolate heterogeneous CTCs. To compensate for the known disadvantages of the positive and negative enrichments, in this study we introduced a two-stage microfluidic chip. The first stage involves a microfluidic magnetic activated cell sorting (μ-MACS) chip to elute white blood cells (WBCs). The second stage involves a geometrically activated surface interaction (GASI) chip for the selective isolation of CTCs. We observed up to 763-fold enrichment in cancer cells spiked into 5 mL of blood sample using the μ-MACS chip at 400 μL/min flow rate. Cancer cells were successfully separated with separation efficiencies ranging from 10.19% to 22.91% based on their EpCAM or HER2 surface protein expression using the GASI chip at a 100 μL/min flow rate. Our two-stage microfluidic chips not only isolated CTCs from blood cells, but also classified heterogeneous CTCs based on their characteristics. Therefore, our chips can contribute to research on CTC heterogeneity of CTCs, and, by extension, personalized cancer treatment. & 2014 Elsevier B.V. All rights reserved.
Keywords: Circulating tumor cell (CTC) Selective isolation Microfluidic magnetic activated cell sorting (micro-MACS) Geometrically activated surface interaction (GASI) Epithelial cell adhesion molecule (EpCAM) Human epidermal growth factor receptor 2 (HER2)
1. Introduction Tumor heterogeneity has been observed between tumors of the same histopathological subtype (inter-tumor heterogeneity) or within a primary tumor and its metastases and even within a single tumor at either a primary or a metastatic tumor site (intratumor heterogeneity) (Hayes and Paoletti, 2013; Fisher et al., 2013). Inter-tumor heterogeneity is well established in breast and colon cancers and other solid tumors (Cusnir and Cavalcante, 2012). According to the characteristics of tumor subtypes, remarkable differences exist in prognosis and therapeutic strategies (Goldhirsch et al., 2011). In the latter case, intra-tumor heterogeneity may be important in personalized medicine approaches where single tumor biopsy samples are generally needed to portray tumor mutational landscapes (Gerlinger et al., 2012). n Corresponding author at: School of Mechanical Engineering, Yonsei University, 262 Seongsan-no, Seodaemun-gu, Seoul 120-749, Republic of Korea. Tel.: þ 82 2 2123 5814; fax: þ82 2 312 2159. E-mail address: [email protected] (H.-I. Jung). 1 These authors contributed equally to this paper.
In this light, circulating tumor cells (CTCs) have been considered a viable and readily accessible alternative source of tumor cells in the form of a “liquid biopsy” to clarify the intra-tumor heterogeneity (Cristofanilli et al., 2004). Nevertheless, CTC research has been challenged by the rarity (1–100 CTCs per 1 mL of peripheral blood from a human cancer patient) and heterogeneity of CTCs. Numerous approaches with microfluidics have been developed to obtain pure CTCs from a large number of normal hematological cells. The two main approaches are positive and negative enrichment. The positive enrichment method specifically isolates the target cells (CTCs) from non-target cells (normal hematological cells) using their specific characteristics, such as size, electric charge and surface protein expression, while the negative enrichment method specifically elutes non-target cells (Hyun and Jung, 2014). Each method has its own advantages and disadvantages. The positive enrichment can separate CTCs with high purity, but can be biased by the selection criteria and may result in a loss of heterogeneous CTCs which fail to meet the desired criteria (Hughes and Mattison, 2012; Kim et al., 2013; Ozkumur et al.,
http://dx.doi.org/10.1016/j.bios.2014.07.019 0956-5663/& 2014 Elsevier B.V. All rights reserved.
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2013). Although negative enrichment can isolate intact and heterogeneous CTCs, the purity is generally lower than that of positive enrichment (Hyun et al., 2013a, 2013b; Ozkumur et al., 2013; Chen et al., 2011). Additionally, because of their rarity, CTCs must be enriched from a large volume of blood (typically 7.5 mL) to obtain reliable information. Microfluidic devices can continuously isolate CTCs with lower cell loss than a batch process and manipulate small amounts (10 6–10 12 L) of fluid. Thus, along with recovery of heterogeneous CTCs and purity, throughput of microfluidic devices is also an important factor for assessing the performance of CTC isolation. Although several microfluidic devices can isolate various CTCs and then verify their heterogeneity such as the expression ratio of epithelial and mesenchymal cell markers, the research only identifies the existence of heterogeneous CTCs without reseparating the cells based on their characteristics (Hyun et al., 2013b; Yu et al., 2013). Herein, we present a two-stage microfluidic chip for selective isolation of CTCs. The two-stages consisted of two types of microfluidic chips; 1) a microfluidic magnetic-activated cell sorter (μ-MACS) which uses planner microfluidic channels and magnetic fields to deplete magnetically labeled cells from the mixed cells in a continuous flow and, consequently, achieves a highly efficient sorting of the target cells and 2) a geometrically activated surface interaction (GASI) chip which exploits enhanced interactions of target cells and the surface inside the microfluidic channel. In the first stage, the magnetic nanoparticle-coated white blood cells (WBCs) were eliminated in the magnetic field whereas nonlabeled CTCs were freely released into the next stage. In the second stage the released CTCs were captured in the GASI chip depending on their expression level of epitope (e.g. EpCAM or Her2). Our experimental results showed the two-stage microfluidic chip is very useful for selectively and specifically isolating CTCs having varying surface protein expression.
2.2. Immobilization of captured molecules in the GASI channel EpCAM or HER2 antibodies were used to capture EpCAM positive cells, or HER2-positive cells inside the channel, respectively. To coat the channel with these antibodies, an affinity surface was prepared using established protocols with slight modifications. First, a solution of 0.2 mg/mL neutravidin in 10 mM Tris buffer was introduced into the channel using a syringe pump (flow rate: 100 μL/min), followed by incubation for 1 min and then rinsing with 10 mM Tris buffer. Then, each biotinylated antibody was added to complete the surface conjugation by forming the avidin–biotin complex and incubated for 10 min, followed by two rounds of rinsing with phosphate buffered saline (PBS) containing 1% bovine serum albumin (BSA) to prevent nonspecific binding. All reagent incubations were performed at room temperature. 2.3. Sample preparation Various human breast cancer cell lines (MCF-7, SK-BR-3, MDAMB-231) were used as the CTC model to evaluate the performance of the chips. The cells were cultured at 37 °C with 5% CO2. After detachment from the plate by trypsin treatment, the cells were resuspended in 10 mL PBS containing 2% fetal bovine serum (FBS). Human blood samples (5 mL) from a healthy volunteer were collected in vacutainer tubes containing EDTA and processed within 6 h. For the elimination of erythrocytes, we used hemolysis buffer (Qiagen, Chatsworth, CA, USA), which was added to the blood samples in a 1:10 v/v ratio and incubated for 10 min at room temperature. The samples were centrifuged at 600g for 10 min and the pellets resuspended in 3 mL of PBS containing 2% FBS. EasySeps human whole blood CD45 depletion cocktail (StemCell Technologies, North America, www.stemcell.com) was added to the samples in a 1:10 blood volume/cocktail volume ratio and incubated for 15 min to coat the CD45 antibody-conjugated magnetic nanoparticles to leukocytes. EasySeps magnetic nanoparticles were added in sequence as described above and the mixture was incubated for 10 min.
2. Experimental section 2.4. Cell staining 2.1. Microfluidic chip fabrication A microfluidic chamber was fabricated using a laser machine (Universal Laser Systems, Scottsdale, AZ, USA) for rapid prototyping. The laser machine cut the designed microfluidic chamber from a PMMA sheet that was sandwiched between double-sided tapes (3M 9475LE; 3M Company, St. Paul, MN, USA). The top and bottom protective layers of the cut PMMA sheet were peeled away revealing the permanent adhesive surfaces. Then, each laser-cut polyester film (3M Hydrophilic Polyester Film 9962; 3M Company, St. Paul, MN, USA) was pressed on the top and bottom of the permanent adhesive surface. The chamber height was 930 mm and total volume 1 mL. To create a magnetic field gradient for efficient attraction of magnetic nanoparticles, NdFeB block magnets (EP331, EP332; e-Magnets UK, Hertfordshire, UK) with a maximum energy product of 45 MGOe (grade N45) were arrayed in a laser-cut plastic cartridge (Fig. S1). A strong magnetic field was generated between the magnetic field lines of two adjacent magnets. The arrangement was parallel to the flow direction and provides magnetic nanoparticles with longer exposure time within the induced magnetic field gradient. The microfluidic chamber was sandwiched between the magnet array cartridges to increase the possibility of capturing leukocytes labeled with magnetic nanoparticles. Conventional two layer soft lithography was used to pattern the focusing þGASI channel (Hyun et al., 2013b).
The cells collected from the outlet were fixed on a glass slide using cytospin and subsequently stained with immunofluorescence for DAPI, cytokeratin, EpCAM and CD45 to identify nuclei, CTCs, epithelial cells and leukocytes, respectively. Finally, the slide glass was fitted on an automated image analysis system and photographed by a moving stage linked with an image-taking system programmed by MATLABs. The cells were labeled with the PKH26 red fluorescent cell linker kit purchased from Sigma-Aldrich according to the manufacturer's protocol with minor modifications. Briefly, the cell suspension was washed with PBS twice (centrifugation was performed at 400g for 5 min) and the cells were suspended in 1 mL of Diluent C (Sigma-Aldrich). Simultaneously, 4 μL of PKH26 dye (Sigma-Aldrich) was added to 1 mL Diluent C and incubated with the cell solution for 5 min at room temperature. To remove the excess dye, 2 mL of 1% BSA in PBS was added and incubated for 1 min. Cells were then resuspended in 10 mL of complete media and washed three times. 2.5. Operation We used a syringe pump (KDS210, KD Scientific Inc., MA, USA) to inject the prepared sample continuously into the channel. The flow rate in the μ-MACS chip was from 100 to 400 μL/min and in the GASI chip from 80 to 100 μL/min. Prior to μ-MACS chip
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experiment, we performed a degassing process by filling the channel with PBS with 2% FBS to prevent non-specific binding of cells. To observe the enrichment efficiency after the process, the cells collected from the outlet were spin-coated on a slide glass at 1500 rpm for 5 min. An image of the entire area was captured using auto-stage coupled fluorescence microscope. Using the target cell identification criteria, the captured images were examined.
3. Results and discussion 3.1. Design and function of the two-stage microfluidic chip Previously, we introduced a GASI chip to isolate intact and heterogeneous CTCs (Hyun et al., 2013b). This technique employs an asymmetric herringbone that generates enhanced mixing flows, increasing the surface interaction between the cells and antibody-conjugated channel surface. However, due to low operational flow rate (20 uL/min), this chip was not applicable for the process of fully drawing blood in a clinical setting, typically 7.5 mL. Before injecting the blood samples into the GASI chip, the cells were preconcentrated into 1–2 mL of buffer by conventional centrifugation to handle the large sample volume, but the capture efficiency with a high concentration of cells decreases because the captured cells tend to detach by interaction with cells traveling in the flow stream (Sin et al., 2005). Additionally, although this chip can isolate heterogeneous
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CTCs, the cells need to be sorted based on their characteristics to investigate their heterogeneity. Therefore, we developed a new two-stage enrichment chip to improve the throughput and selectively isolate the CTCs based on their surface marker. In the first stage (μ-MACS), WBCs coated with CD45 antibody-conjugated magnetic nanoparticles were captured inside the chip by magnetic force and heterogeneous CTCs exited through the outlet. After the first-stage of enrichment, all the cells were congregated in the center of the channel by inertial force and then moved into the GASI region. When the cells entered the GASI region, the EpCAM-positive cells were specifically captured on the EpCAM antibody-coated channels, while the EpCAM-negative cells were eluted through the outlet (Fig. 1). Currently, the conventional batch MACS system is too bulky in size to carry for point-of-care medical systems (Chen et al., 2014) and cannot provide a continuous process necessary to eliminate the risk of losing target cells (Hyun and Jung, 2014). Therefore, microfluidic technology is needed to provide portability and continuous flow for more precise control of the magnetic and hydrodynamic forces acting on the target cells to improve separation. The μ-MACS chip consists of an inlet, an outlet, four seriallyconnected straight channels, and two neodymium magnetic plates. Because the magnetic flux decreases dramatically with distance from the magnet surface (Pamme, 2006), thin polyester films (1-mm-thick) were used as fluidic channel covers, which bordered both the top and the underside of the magnet to increase the bias of magnetic nanoparticle-tagged cells to the magnetic force. The magnet plate was composed of 21 rectangular
Fig. 1. Schematic diagram of a two-stage enrichment chip for selective isolation of CTCs. The two-stage microfluidic chips consisted of a microfluidic magnetic activated cell sorter (μ-MACS) and an inertial focusing channel combined with a geometrically activated surface interaction (GASI) chip. After the μ-MACS chip eliminated a number of white blood cells (WBCs) via negative enrichment, the GASI chip was used for positive enrichment to selectively isolate cells based on their surface protein expression.
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neodymium magnets. The two neodymium magnetic plates were assembled to the microfluidic channel by sandwiching. Although 99% of WBCs were eliminated with the first-stage μMACS chip, a large sample volume remains a throughput issue for the GASI chip. Therefore, we devised a strategy to enhance the throughput of the GASI chip by adding a focusing channel. Prior to entering the GASI region, the sample flow was decelerated by the inertial focusing channel. In a rectangular microchannel, the shearinduced lift force drives particles toward the channel walls while the wall-induced lift force drives particles toward the channel centerline. The microparticles migrate to the equilibrium position by the balance of the two lift forces. In particular, particles focus into two equilibrium positions centered at the top and bottom walls at a low aspect ratio (AR ¼height/width) microchannel at a moderate Reynolds number (Re), which is expressed by the following equation:
Re =
ρUmDh μ
where ρ is the density of the fluid, Um is the maximum flow velocity in the channel, Dh is the hydraulic diameter of the channel, and μ is the dynamic viscosity of the fluid. Reynolds number is a dimension less number that prioritizes between inertial force and viscous force in the microchannel (Zhou and Papautsky, 2013). Thus, the focusing channel was designed to draw all the cells into the GASI region and discard a portion of the buffer to the outlets as waste consequently, the GASI region adjusted by changing channel dimensions. The flow rate was directly proportional to the pressure drop in the microchannel because it was the product of the cross-sectional area of the channel and average velocity. As mentioned above, because the GASI operated at a low flow rate, we determined the channel dimensions for the ratio of waste outlet flow rate to the GASI region to be nearly 7:3. The height and width of focusing channels were 40 and 100 μm, respectively (AR ¼0.4) and the focusing channel length was 14 cm. The width of waste channels was 150 μm, while the width of channel toward the GASI region was 200 μm. 3.2. Determination of experimental conditions and evaluation of chip performance We determined the experimental conditions for the first-stage negative enrichment process (μ-MACS chip). To investigate the depletion limit of the μ-MACS chip, we prepared three different volumes of hemolyzed blood samples from healthy donors (3, 5,
and 7.5 mL) resuspended in 3 mL of PBS with 2% FBS and then conjugated with magnetic nanoparticles. We loaded the samples into the inlet and then counted the number of cells that exited the outlet using a hemocytometer. The depletion efficiency was over 99% regardless of cell concentration and flow rate because the magnetic flux density of the neodymium magnets was sufficient for capturing the magnetic nanoparticle-coated leukocytes (almost 1700 G per neodymium magnet) (Fig. 2a). In CTC research, recovering as many target CTCs as possible is also important due to their rarity. When the leukocytes and cancer cells were loaded at a fixed concentration of 1 × 106/mL into the inlet, 90.97% of cells were recovered at 400 μL/min inlet flow rate (Fig. 2b). Although the high flow rate dictates the high throughput performance of a chip, sample leakage was observed at the upper flow rate (400 μL/min) because of the high pressure. Thus, we determined the optimal flow rate as 400 μL/min. The number of cancer cells (MCF-7) isolated using the μ-MACS chip was proportional to the number of cancer cells spiked into blood samples across the physiologically relevant range (50–1000 cancer cells/5 mL of human blood) (Fig. 2c). The enrichment yield was as high as 763.14 fold. In the second stage (GASI chip), the enriched cells on the μMACS chip were reseparated based on their surface markers by capture with specific antibody-coated GASI channels. Regarding a recent study reporting that CTCs and WBCs overlap in size (Kim et al., 2012), we considered all the cells as CTCs because the purpose of the second stage enrichment was reclassification of cells, not elimination of leukocytes. Thus, the green 7- and red 15μm fluorescent beads were tested at a fixed concentration of 1 105/mL to ensure all the cells congregated into the center of the channel. Although the focusing efficiency increased as the Re increased in the low aspect ratio microchannel, increasing the flow rate indefinitely was difficult because the high shear force militated against affinity-based cell capture in the GASI region. Therefore, we designed a sufficiently long focusing channel to concentrate the small microbeads into the channel center at a moderate Re (Re¼38–47). Within the focusing channel, relatively larger 15μm beads were well focused across all 80–100 μL/min flow rates (from 97.3571.98% to 98.23 73.54%; n ¼3). Regarding the 7-μm beads, the focusing efficiency increased from 87.18 74.17% to 95.01 71.95% as the flow rate increased (Fig. 3a). Therefore, we determined the optimal flow rate to be 100 μL/min. Fig. 3b shows the behavior of fluorescent microbeads in the focusing channel and outlets at the 100 μL/min flow rate. When three different concentrations of hemolyzed blood samples and cancer cell
Fig. 2. Determination of optimal conditions. (a) Depletion efficiency depending on the flow rate and concentration of blood cells. The depletion efficiency was almost 100% with these terms. (b) Effects of flow rate on recovery. The recovery was under 10% at a flow rate from 100 μL/min to 200 μL/min and increased to approximately 90% at a flow rate of 300 μL/min. (c) Regression analysis of capture yield for a range of MCF-7 cell concentrations spiked into whole blood. The number of separated cells was proportional to the increased number of spiked cells (y = 0.913x + 8.3888, R2 = 0.9991) .
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Fig. 3. (a) Focusing efficiency based on flow rate and bead size. Flow rates from 80 to 100 mL/min did not cause any significant differences in efficiency. However, the focusing efficiency with the 15-μm bead was higher than with the 7-μm bead. (b) Fluorescent images illustrating migration of 7-μm and 15-μm beads toward the microchannel center. Flow is from left to right. (c) Focusing efficiency based on the number of blood cells.
samples were injected individually into the inlet, the 100-fold diluted blood sample reached almost 100% focusing efficiency and 95.39% of cancer cells were successfully focused (Fig. 3c). While the focusing efficiency was relatively lower with undiluted than diluted blood samples, this was an acceptable value because a large amount of blood cells were eliminated by first-stage negative enrichment. 3.3. Selective separation of cancer cells To verify that the GASI chip selectively separated the CTCs based on their surface markers, we evaluated the separation yield using three different breast cancer cell lines (MCF-7, SK-BR-3 and MDA-MB-231 cells). EpCAM surface densities on the cell lines ranged from 2.22 105 binding sites/cell on MCF-7 to 1.7 103 binding sites/cell on MDA-MB-231. The HER2 surface densities ranged from 9.76 105 binding sites/cell on SK-BR-3 to 14.7 103 binding sites/cell on MDA-MB-231 (Prang et al., 2005). The results indicated the MCF-7, SK-BR-3 and MDA-MB-231 cells can be considered as EpCAM-positive, HER2-positive, and both EpCAMand HER2-negative, respectively. We successfully separated tumor cells based on their EpCAM and HER2 expression levels on the anti-EpCAM antibody-coated chip and anti-HER2 antibody-coated chip, respectively (98.81% of MCF-7 cells (EpCAMþ ) and 93.12% of MDA-MB-231 cells (EpCAM ), 86.51% of SK-BR-3 cells (HER2 þ) and 66.54% of MDA-MB-231 (HER2 )) (Fig. 4). It was thought that the reason why the EpCAM based separation yield is higher than that of HER2 based separation is that the difference of EpCAM expression level between MCF-7 and MDA-MB-231 cells (approximately 130 times) is greater than the difference of HER2
expression level between SK-BR-3 and MDA-MB-231 cells (roughly 66 times). In clinical samples, the heterogeneity of cells and their fractions remain undetermined; therefore, mixed cell experiments were conducted to show the GASI chip can handle any fraction of samples. When the mixed MCF-7 cells and MDA-MB-231 cells were injected into the anti-EpCAM antibody-coated GASI chip, the cells were sorted according to their surface markers, even if the smaller ratio of MCF-7 and MDA-MB-231 cells induced a lower purity in MCF-7 cells similar to the inverse ratio. The mixed cell experiments with SK-BR-3 and MDA-MB-231 cell lines produced similar results (Table 1). Finally, we tested the two-stage enrichment by separating cancer cell lines spiked into 5 mL of blood samples. Sample 1 was prepared by spiking 103 of MCF-7 cells (green cell tracker staining) and 103 of MDA-MB-231 cells (red cell tracker staining) per 1 mL of blood and sample 2 was prepared from 103 of SK-BR-3 cells (green cell tracker staining) and 103 of MDA-MB-231 cells (red cell tracker staining) at the same concentration as sample 1. The WBCs in these two samples were eliminated by the μ-MACS chip and then sample 1 was injected into the EpCAM antibodycoated GASI chip while sample 2 was infused into the HER2 antibody-coated GASI chip. After the samples passed through the GASI chip, the separated cells were fixed on a slide glass and stained with DAPI for DNA content. After imaging the slides, the captured images were examined according to predefined criteria (MCF-7 and SK-BR-3 cells: DAPI positive, green fluorescent positive; MDA-MB-231 cells: DAPI positive, red fluorescent positive; WBCs: DAPI positive, green and red fluorescent negative) (Fig. 5). Because the GASI region is too large to take pictures, while the
Fig. 4. Second-stage (GASI chip) enrichment for selective capture of CTCs. Tumor cells were selectively separated based on their EpCAM and HER2 expression levels on a GASI chip (98.81% of MCF-7 cells (EpCAM þ ) and 93.12% of MDA-MB-231 cells (EpCAM ), 86.51% of SK-BR-3 cells (HER2 þ) and 66.54% of MDA-MB-231 cells (HER2 )).
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Table 1 Comparison of purity according to the ratio of mixed cancer cells. Ratio (MCF-7/MDA-MB231) [SK-BR-3/MDA-MB231]
MCF7 Purity (%) [SK-BR-3 Purity]
MDAMB231 Purity (%) [MDA-MB-231 Purity]
1:1 1:10 1:100 10:1 100:1
50.63 [55.67] 9.61 [9.80] 0.24 [1.04] 91.10 [91.09] 99.43 [99.14]
80.39 [71.98] 96.85 [94.66] 91.10 [99.24] 11.04 [10.35] 5.22 [1.90]
cells passed through the outlet is spin-coated in 1 mm diameter circle area, the MCF-7 and SK-BR-3 cells (yellow dashed circles) appeared fewer than MDA-MB-231 cells (white dashed circles). The purities used for measuring the purity of the target cells were 22.91% for MCF-7 and 15.43% for MDA-MB-231 in EpCAM-based selection, 13.37% for SK-BR-3 and 10.19% for MDA-MB-231 cells in HER2-based selection. Although a small amount of WBCs remained, the purity of the spiked cancer cells increased significantly to 22.91% in the two-stage microfluidic chip (μ-MACS chip and GASI chip) compared with 3.64% in the GASI chip.
Additionally, our two-stage microfluidic chip can be used as two-stage negative enrichment to produce high enrichment yields. When the cells were injected into the first stage μ-MACS chip and then second stage GASI chip, the negative enrichment ratio increased by approximately 44 106 fold (data not shown).
4. Conclusion CTCs have been certified as diagnostic and prognostic indicators of cancer in many clinical studies with the help of the many techniques that have been introduced for CTC isolation. However, the heterogeneity of CTCs, which may be important in personalized cancer medicine and for the development of new anti-cancer drugs has not been determined. In this study, we constructed a high-throughput and high-purity two-stage microfluidic chip for WBC elimination and selective CTC isolation (Table S1). The first stage (μ-MACS) was a high throughput negative enrichment for WBCs elimination via magnetophoresis. In the second stage (specific antibody-coated GASI chip), the cells collected from the μ-MACS chip were selectively separated based on the affinity of specific antibodies for their surface proteins. Because antibodies
Fig. 5. Fluorescent images of cells selectively isolated using a two-stage microfluidic chip. Each cell was identified by predefined criteria (MCF-7 and SK-BR-3 cells: DAPI positive, green fluorescent positive; MDA-MB-231 cells: DAPI positive, red fluorescent positive; WBCs: DAPI positive, green and red fluorescent negative) (For interpretation of the references to color in this figure, the reader is referred to the web version of this article).
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can be selected based on the research purpose, our two-stage microfluidic chip was designed to provide various opportunities for cell heterogeneity investigation in rare cell research including CTCs. When the 3-mL sample was used, the entire process from the μ-MACS chip to the GASI chip was completed in 1 h.
Acknowledgments This study was supported in part by a grant from the Korea Health Technology R&D Project (Grant no. A121986) of the Ministry of Health & Welfare and the Center for BioNano Health-Guard funded by the Ministry of Science, ICT & Future Planning (MSIP) of Korea as Global Frontier Project (H-GUARD_2013M3A6B2078959).
Appendix A. Supplementary information Supplementary data associated with this article can be found in the online version at http://dx.doi.org/10.1016/j.bios.2014.07.019.
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