Biomedical Microdevices

, Volume 11, Issue 4, pp 903–913

A microfluidic cell culture platform for real-time cellular imaging


  • Chia-Chun Hsieh
    • Department of Engineering ScienceNational Cheng Kung University
  • Song-Bin Huang
    • Department of Engineering ScienceNational Cheng Kung University
  • Ping-Ching Wu
    • Institute of Basic Medical SciencesNational Cheng Kung University
    • Institute of Oral Medicine and Department of Stomatology and Institute of Innovation and Advanced StudiesNational Cheng Kung University
    • Department of Engineering ScienceNational Cheng Kung University
    • Medical Electronics and Device Technology CenterIndustrial Technology Research Institute

DOI: 10.1007/s10544-009-9307-7

Cite this article as:
Hsieh, C., Huang, S., Wu, P. et al. Biomed Microdevices (2009) 11: 903. doi:10.1007/s10544-009-9307-7


This study reports a new microfluidic cell culture platform for real-time, in vitro microscopic observation and evaluation of cellular functions. Microheaters, a micro temperature sensor, and micropumps are integrated into the system to achieve a self-contained, perfusion-based, cell culture microenvironment. The key feature of the platform includes a unique, ultra-thin, culture chamber with a depth of 180 μm, allowing for real-time, high-resolution cellular imaging by combining bright field and fluorescent optics to visualize nanoparticle-cell/organelle interactions. The cell plating, culturing, harvesting and replenishing processes are performed automatically. The developed platform also enables drug screening and real-time, in situ investigation of the cellular and sub-cellular delivery process of nano vectors. The mitotic activity and the interaction between cells and the nano drug carriers (conjugated quantum dots-epirubicin) are successfully monitored in this device. This developed system could be a promising platform for a wide variety of applications such as high-throughput, cell-based studies and as a diagnostic cellular imaging system.


MicrofluidicsMicro-bioreactorsCell cultureNanoparticleCellular imagingMEMS





Charge-coupled device


Depth of field


Electromagnetic valve


Fetal bovine serum


4-(2-hydroxyethyl)-1-piperazineethanesulfonic Acid




3-(4,5-Dimethyl-2-thiazolyl)-2,5-diphenyl-2H-tetrazolium Bromide


Oral cancer cell






Quantum dots



1 Introduction

Multi-modal microscopic imaging has long been a crucial technique for the study of cell biology and molecular biology. Microscopy techniques have been applied for assessing cell population dynamics, single-cell or even sub-cellular structures and functions, and cellular response to defined treatments. Moreover, cell-based assays for evaluating efficacy (Lang et al. 2006; Partlow et al. 2008; Gupta et al. 2007), biological toxicity (Xu et al. 2008; Earl et al. 2008), therapeutic modalities and nanoparticle-based formulations (Tseng et al. 2007; Patra et al. 2007; Keter et al. 2008; Wu et al. 2008a; Wu et al. 2007a) greatly rely on a full understanding of the cellular responses to the treatment and its micro-environments. For example, porous iron-oxide nanorods have been developed as delivery nanocapsules (Wu et al. 2007a). Their drug delivery efficiency, sub-cellular targeting, and therapeutic efficacy have been confirmed from microscopic observation of the morphological alterations and the viability of the treated Hela cells. Besides, dynamic time-lapsed, long-term observation of cell-based assays plays important roles in the understanding of the cell and in molecular biology but are technically challenging, especially when multiplexed, high-throughput observation are desired. In order to get high-throughput, in-depth insights from the interaction between the drug delivery system and the test cells, an automatic cell culture platform capable of microscopic imaging and precise quantification of efficacy under an in vivo-like microenvironment is crucially in demand.

Microsystems based on microfluidic control technology have been demonstrated to be suitable for cellular micro-assays. For example, a microchip device has been demonstrated to improve the cell growth of primary hepatocytes (Powers et al. 2002). Microfluidic-based devices for long-term cell culture and imaging have been extensively studied recently (Kaji et al. 2003; Sin et al. 2004) such as a cell culture chip for on-line monitoring of cultured eukaryote cells (Michael et al. 2006) and the control of axon growth and polarization from a central nervous system using a microfluidic culture platform without modulation of neurotrophins (Taylor et al. 2005). Elastomeric microchannel systems could even support in vitro mouse embryo development (Beebe et al. 2002; Raty et al. 2004). However, most of these cell culture devices require large equipment such as incubators and syringe pumps to maintain cell culture environments. These apparatus are usually costly and bulky, thus hindering their practical applications. We recently reported an automatic cell culture chip integrating a micro temperature control module and medium transport mechanisms in a single chip (Huang et al. 2007). However, the flow rates from the micropumps may influence cell physiology due to the induced high shear stress. Thus a non-continuous medium supply system which controlled the shear stress level in the culture medium was implemented. Moreover, this advanced microfluidic culture system not only kept the culture system sterile, providing an adequate nutrient supply and waste removal during the culture period, but also maintained a stable, defined and flexible culture microenvironment for delivery of externally applied stimuli such as drugs or light (Wu et al. 2006; Sittinger et al. 1997; Wu et al. 2007b). A perfusion-based, micro, cell culture chip using a spider-web micropump has been developed to provide uniform and simultaneous culture medium delivery and replacement in 8 microchannels (Wu et al. 2008b). Although the design of the spider-web shape micropumps are suitable for simultaneous multiple pumping, the flow rate in the microchannels is fixed and thus does not allow tunable shear stress inside each channel. Hence, in order to solve this problem and to provide an adequate flow rate for continuous medium supply, a membrane-based serpentine-shape (S-shape) pneumatic micropump has been proposed (Huang et al. 2008). This is achieved by fine-tuning the fluidic resistance of injected air in the designed pneumatic microchannels. Besides, the S-shape pneumatic micropump can be integrated into a perfusion-based cell culture platform to continuously supply medium (Wu et al. 2008c). In this study, for high-magnification microscopic observation of the cultured cells and their response to drug delivery, we designed an ultra-thin window to be compatible with an objective lens with a depth of field (DOF) of as small as 200 µm, in order to achieve adequate spatial resolution. The new cell culture platform has a modified S-shape micropump with micro check valves to achieve better control of the flow rate. This enables a continuous medium supply at a lower flow rate for general cell culture and at a higher flow rate for efficient cell loading into the micro-culture area; both of which can now achieved in the same device. In addition, cross-contamination is prevented by the micro check valve design. The ring-shape pillars with gradually increasing spacing intervals encircling the cell culture area could effectively reduce the shear stress induced by the medium flow. With this integrated approach, the developed system provides a new platform capable of potential high-throughput, high-magnification observation for real-time, cell-based study and functional assays.

2 Experimental

2.1 Chip design

The layout of the proposed cell culture platform for real-time, high-magnification, cellular imaging is shown in Fig. 1(a). The platform comprises two modules, including a cell culture module for continuous medium supply and for cellular imaging, and a micro temperature control module for maintaining the cell culture environment at 37°C. The cell culture module is comprised of five S-shape pneumatic micropumps (Huang et al. 2008) for providing a continuous medium supply and cell loading, five micro check valves which are designed to keep a three-dimensional block inside the channel (Huang et al. 2007) to prevent cross contamination, microchannels for cell loading, medium supply and drug delivery, six reservoirs for storage of test drugs, fresh medium and waste fluid, and an ultra-thin, micro-culture chamber. The detailed pumping mechanism of the integrated S-shape pneumatic micropump (Huang et al. 2008) connected to a micro check valve is illustrated in Fig. 1(b). The operation starts with loading of the fresh medium into the “reservoirs for medium or drugs” and stopped by the micro check valve. Based on the fluidic resistance of air in the S-shape pneumatic channel, the sequential raise of air pressure in the pneumatic channel forms a peristaltic sequence of three PDMS membrane deflections due to the injection of compressed air. Successive deformation of the PDMS membranes moves the medium forward. At the same time, the hydrodynamic force generated by the micropump lifts the block structure of the micro check valve and enables medium flow through the next chamber. Thus the sequential release of the PDMS membranes ultimately re-fills the channel with the medium from the reservoirs. Besides, the medium, being pumped forward through the micro check valve, is blocked from flowing in the reversed direction since the microchannel is immediately in the close state after the medium flows past the valve. With this approach, the chance of cross contamination between samples is minimized. Notably, two different pumping directions are designed into the cell culture module, one for cell loading and another for medium/drug loading. For the cell loading process, the suspended cells are first loaded into the “reservoir for waste and cell loading” (Fig. 1(a)). The first S-shape pneumatic micropump is used to provide a suction force which allows the cells to flow into the cell culture area. This approach prevents the cells from being crushed against the micro check valve. For the drug loading process, the drugs dissolved in the medium are first loaded into the respective “reservoir for medium or drugs” (Fig. 1(a)) and then are delivered into the cell culture area at a controlled flow rate, in sequence, by using each individual micropump. With this approach, the cells can be tested continuously with up to four different drugs at the same time in the current design. Figure 1(c-I) presents an exploded view of the cell culture module comprising of three layers including an air chamber layer, a fluidic channel and cell culture chamber layer, and a bottom glass layer, which are permanently bonded together to form a laminate structure encompassing channels and chambers (Figs. 1(c-I) and 1(c-II)). As shown in Fig. 1(c-II), one of the key features of the cell culture module is an ultra-thin, cell culture area with a depth of 180 μm, allowing for high-magnification, real-time cellular imaging via microscopic observation using a 100X objective oil-lens. Furthermore, a top-view layout of the cell culture area is illustrated in Fig. 1(d-I). The expanding channel between the cell culture area and the microchannel is designed in a funnel shape to allow the microchannel (500 μm wide initially) to gradually extend to the center of the culture area (5,000 μm wide). Due to conservation of mass flow rate, this design also slows down the medium flow rate from the microchannel inlet into the cell culture area such that the influence of the flow shear stress can be minimized. Moreover, the ring-shape pillars with gradually increasing pillar lengths between fixed gaps (with pillar lengths of 40 μm, 50 μm, 60 μm, 70 μm, 80 μm, 90 μm, 100 μm, 110 μm, and 200 μm sequentially) are designed to encompass the cell culture area and to further reduce the shear stress induced by the medium flow. In addition, a streamline-shape structure (center pillar) located in the center of the culture area is designed to prevent the top membrane from collapsing.
Fig. 1

(a) A schematic illustration of the cell culture platform. It is comprised of two modules, including a cell culture module and a micro temperature control module. (b) Schematic illustration of an S-shape micropump with a micro check valve. (c) An exploded view of the cell culture module (I), and a cross-sectional view of the cell culture module (II). (d) A close-up view of the cell culture area equipped with a ring-shape array of pillars encircling the cell culture area in order to reduce the shear stress induced by the culture medium flow (I), and a close-up view of the micro temperature controller chip (II)

The micro temperature control module is composed of multiple microheaters and a micro temperature sensor (Fig. 1(d-II)) to maintain a constant temperature suitable for the cell culture process. Detailed information about the microheaters and the micro temperature sensor can be found in our previous work (Hsieh et al. 2006).

2.2 Fabrication

Two major fabrication processes are involved in creating the cell culture platform. For the cell culture module, the fabrication processes are based on SU-8 lithography and polydimethylsiloxane (PDMS, Sylgard® 184, Dow Corning Inc., USA) replication processes. The detailed fabrication processes can be found in our previous work (Huang et al. 2008). Briefly, for the master template for the S-shape micropumps, a layer of SU-8 25 negative photoresist (MicroChem, USA) with a thickness of 25 μm is first patterned on a silicon wafer (Fig. 2(a-I), (a-IV)) to form the structures of the pneumatic microchannels. Then, another layer of SU-8 50 negative photoresist (MicroChem, USA) with a thickness of 100 μm is patterned on the same silicon wafer to form air chambers (Fig. 2(a-V)). Similarly, for the master template of the microchannel and the cell culture area, a layer of SU-8 50 photoresist with a thickness of 100 µm is patterned on another silicon wafer (Fig. 2(a-VI)). Once these two master templates are finished, the inverse images of the microstructures are fabricated by using a PDMS replication process. PDMS polymer is prepared by thoroughly mixing the PDMS pre-polymer and curing agent in a ratio of 10:1 by weight. The polymer is then de-aerated under vacuum to remove all air bubbles created during mixing. The mixture is then poured onto the master templates and cured at 70°C for 2 h. Note that a PDMS membrane with a thickness of 80 μm is formed by controlling the spin-coating speed of PDMS. All of the cured PDMS layers are then obtained after a careful mechanical de-molding process. Two layers of PDMS and the bottom glass (thickness =130 μm) are then bonded by using an oxygen plasma treatment to form a cell culture platform (Fig. 2(a-VII)).
Fig. 2

Schematic illustration of the simplified fabrication process for the cell culture module (a) and the micro temperature control module (b). (a-I ~ a-V): a lithography process is used to define a SU-8 template for the S-shape micropumps. (a-VI): a lithography process is used to define a SU-8 template for the microchannel and the cell culture area. (a-VII): After the PDMS replication process to fabricate inverse images of the SU-8 microstructures, oxygen plasma treatment is performed prior to bonding the PDMS layers and the cover glass substrate to form a complete chip. (b-I ~ b-III): an ITO layer is etched to form the microheaters. (b-IV): a layer of platinum (Pt, 0.07 μm) for the temperature sensor is deposited and patterned by using an electron-beam evaporation and a standard lift-off process. (b-V): a layer of gold (Au, 0.12 μm) for the electrodes is deposited and patterned

For the micro temperature control module, indium-tin-oxide (ITO, 7Ω/□, Ritek Corp., Taiwan) glass is used for real-time cellular imaging. This detailed fabrication process can be also found in our previous paper (Huang et al. 2007). Briefly, two ITO-based microheaters are fabricated by using AZ4620 positive photoresist (Clariant Corp., Switzerland) and an ITO etching process (Figs. 2(b-I)-(b-III)). Another layer of AZ4620 positive photoresist is spin-coated, followed by a standard photolithography processes. A layer of platinum (Pt, 0.07 μm) is patterned for the micro temperature sensor by using electron-beam evaporation and a standard lift-off process (as shown in Fig. 2(b-IV)). Similarly, the same metallization process is used to form gold (Au, 0.15 μm) electrodes (as shown in Fig. 2(b-V)). After stabilizing the resistance of the sensor by a sintering process, the resistances of the microheater and the micro temperature sensor are measured to be 54.0 Ω and 414.3 Ω, respectively. Finally, the cell culture module is placed on top of the micro temperature control module to form the cell culture platform as shown in Fig. 3. The dimensions of the assembled platform are measured to be 8 cm x 8 cm. Note that mineral oil is added between the two modules to decrease the thermal resistance and to increase the temperature uniformity of the cell culture area.
Fig. 3

A photograph of the cell culture platform. The width and length of the cell culture platform are 8 cm × 8 cm, respectively

2.3 Experimental setup

The entire experimental setup is schematically illustrated in Fig. 4. A custom-made hand-held controller which integrates an air compressor (MDR2-1A/11, Jun-Air Inc., Japan), five electro-magnetic valves (EMV, S070M-5BG-32, SMC Inc., Taiwan) and a programmable control circuit system is used to modulate the five on-chip micropumps for medium supply and drug delivery. A temperature control system providing the feedback signal from the micro temperature sensor to regulate the microheaters provides a constant micro-culture temperature at 37°C (Hsieh et al. 2006). For high-magnification, real-time cellular and sub-cellular imaging of the micro-culture area, an image capture system comprising of a fluorescent microscope (TE300, Nikon, NY, USA) coupled with a video camera (Camdio, CA-8M3N) is used. The flow rates of the micropumps are characterized by measuring the volume of liquid pumped over a period of 1 h. For these measurements, the pulsation frequency of the PDMS membranes is regulated by the EMV modulated by the custom-made control circuit.
Fig. 4

Block diagram of the experimental components of the cell culture platform. It is composed of a temperature control system, a microchip system, an external air pressure control system and an image recording system

2.4 Sample preparation

In order to evaluate the function of the cell micro-culture platform and the real-time cellular imaging, an oral cancer cell line OC-2 (Huang et al. 2002) is used in this study. The culture medium (HyQ® RPMI-1640; Hyclone, USA) containing 2.05 mM L-glutamine, 10% fetal bovine serum (FBS) (Gibco® 26140-079 US origin, Invitrogen Ltd., Taiwan), 100 units/ml of penicillin and streptomycin (Gibco® 15140-122, Invitrogen Ltd., Taiwan), and 25 mM HEPES (SH30237, HyClone, US) is prepared. The micro-culture platform is sterilized using 70% of ethanol and exposed to ultraviolet light in a laminar flow hood for 30 min. OC-2 cancer cell suspension at 2 × 105 cells/ml density is pumped into the micro-culture area to seed the cells using a high pulsation frequency of 18.20 Hz (flow rate = 1,837.0 µl/hr). After adhesion of the loaded cells (4 h after seeding), the medium is continuously flowed through the cell culture area at a low pulsation frequency of 0.12 Hz to achieve the desired flow rate of 9 μl/hr while the culture environment is kept at 37°C. All reservoirs are covered by sterile circular lids during the culture process to prevent contamination.

In order to observe the influence of “nano drugs” on the cancer cells, we use quantum dots (QDs) conjugated with the anti-cancer drug epirubicin (C27H29NO11, Pfizer, USA) as a model (Olinsji et al. 1997; Wu et al. 2003; Pieper et al. 2000). The quantum dots provide fluorescent labels for imaging the nano drug complex. The epirubicin is conjugated to the QDs (Qdot® 655 ITK, Invitrogen Ltd., Taiwan) with a carboxyl group (-COOH, 1,500 carboxyl/dot) through an amino group (-NH2, 1 amino/molecule) on the molecule by covalent crosslink. The ratio of the anti-cancer drug to conjugate with the QDs is 1,500(epirubicin):1(QDs). A series of cytotoxicity testes are conducted and verified for the QDs, pure epirubicin and the nano drugs using a dimythylthiazol diphenyl tetrazolium bromide (MTT) assay (Fraga et al. 2008) at different treatment dosages (0M, 10−11M, 10−10M, 10−9M, 10−8M for the QDs and the nano drugs and 0M, 1.5 × 10−8M, 1.5 × 10−7M, 1.5 × 10−6M, 1.5 × 10−5M for the epirubicin). The MTT assay is conducted as described previously. Briefly, OC-2 cells are seeded into a 96-well cell culture plate at a final concentration of 5,000 cells/well. After incubation of the seeded OC-2 cells for 24 h, the medium is replaced by fresh culture medium containing different concentrations of QDs and the nano drugs. All experimental conditions are repeated for five times in this study. After incubation for 48 h, 20 µl of MTT solution is added to each well. After incubation for 4 h at 37°C, 100 µl of SDS solution (10% sodium dodecyl sulfate dissolved in 0.01 N HCl) is added to each well. After overnight (18–20 h) incubation, absorbance at 560 nm is measured in an ELISA plate reader (LP 400 Pasteur Diagnostics, France). For the real-time, high-magnification observation of the cells in response to the nano drugs (conjugated QDs-epirubicin), the cell culture medium containing epirubicin conjugated QDs (concentration = 10−8M) is loaded into the reservoirs (see Fig. 1)(a) and a time-lapsed videomicroscopy of the micro-cultured cells is performed continuously for 24 h.

3 Results and discussion

3.1 Characterization of microfluidic components

In order to maintain a continuous supply of culture medium to the cell culture area and to minimize the shear force on the cultured cells, an S-shape micropump with a micro check valve is adopted to provide a low pumping rate. The S-shape micropump is designed to provide better control over a broader flow rate range. This allows a continuous supply of medium at a lower flow rate to reduce the shear stress on the cultured cells while also being capable of loading and replenish cells to the micro-culture area using higher flow rates. Figure 5(a) shows the measured flow rate of a single micropump at various membrane pulsation frequencies at an applied air pressure of 20 psi. The flow rate can be precisely controlled within a range of 10.3 ~ 1,837.0 µl/hr with a saturated flow rate at a pulsation frequency of 18 Hz. This is attributed to the lag in response of the PDMS membranes to the high-frequency air pressure pulsation when the driving frequency is higher than 18 Hz. Figure 5(b) shows the flow rate profile for the micropump operating at lower driving frequencies. At a low pulsation frequency of 0.12 Hz, a flow rate of 9 μl/hr is chosen to deliver the medium through the cell culture area.
Fig. 5

(a) The flow rate profile at different driving frequencies of an S-shape pneumatic micropump at an applied air pressure of 20 psi. (b) The flow rate at lower driving frequencies

Even though the low flow rate range is more suitable for continuous delivery of the medium into the cell culture area, the pulsation produced by the micropumps at the low frequency still may induce excessive shear stress during the cell culture process. Hence, we have designed ring-shape pillars arrayed with gradually increasing pillar lengths to encircle the culture area (Fig. 1)(c) to reduce the shear stress from the medium flow as previously described (Hung et al. 2005). Figure 6 shows an image acquired by a high-speed charge-coupled device (CCD camera (MC1311, Mikrotron, Germany) when red dye is pumped into the cell culture area. The high-speed CCD images show effective reduction of the shear stress by the special layout of the ring-shape pillar array. It is observed that the dye is gently transported into the culture area through gaps between the pillars without flushing the cells away from the seeded culture area at a flow rate below 519 μl/ hr (driving frequency = 4.3 Hz, Pressure = 20psi).
Fig. 6

CCD images of red dye flowing into the cell culture area at a flow rate of 519 μl/hr. it is clearly seen that the dye is transported in a laminar manner into the culture area through gaps between the pillars

A micro temperature control system is implemented in this system. A thermocouple (DE-3003, DER EE, Taiwan) is used to directly measure the temperature variation in the cell culture area. Figure 7 shows that the temperature is maintained at 37°C with a variation of 0.2°C for a measurement period of 28 h. These results indicate that this cell culture platform provides a stable culture temperature micro-environment.
Fig. 7

The temperature measured inside the culture area during a 28-hour period of measurement modulated by a micro temperature control module. A rapid rise in the chamber temperature and a stable temperature (±0.2°C) can be observed

3.2 Cell culture and real-time cell division imaging

As mentioned above, the developed cell culture platform is designed to provide a stable and suitable cell culture micro-environment for the cell harvesting and replenishing process. The micro-culture system is evaluated using an OC-2 cancer cell line. The OC-2 cells are harvested by trypsinization and infused into the system for a continuous culture and recorded by time-lapsed imaging over a period of 72 h. Figure 8(a) shows that the OC-2 cells are seeded initially in a low density (Fig. 8(a-I)) with additional space to allow subsequent cell growth and expansion. After incubating for 48 h, the observation area is almost filled with cells (Fig. 8(a-II)). Moreover, after 72 h, some proliferating cells adhering upon others cells are observed (Fig. 8(a-III)) due to lost of contact inhibition in cancer cells. Figure 8(a) shows the morphology of the cultured oral cancer cells with a normal polygonal shape. Under a 20X objective lens, the dynamic cell division process is observed in real-time (Fig. 8(b)). The cytokinesis process of cell A (from Fig. 8(b-II) to Fig. 8(b- VI)) and cell B (from Fig. 8(b-III) to Fig. 8(b- VI)) after seeding for 18 to 22 h is clearly observed. The cell A first rounds up from the surface of the cell culture area (Fig. 8(b-II)), starts to divide (Fig. 8(b-III)), and then separates into two independent cells (Fig. 8(b-IV)). Finally, the divided cells start to adhere and to flatten down on the cell culture area (Fig. 8(b-VI)). These results indicate that a stable, long-term culture environment is established in the developed cell micro-culture platform.
Fig. 8

(a) Bright-field images of the cancer cells growing inside the cell culture platform. A significant increase in the cell population is observed. The images (I), (II), (III) are captured by a video camera coupled with a microscope at 4, 48, and 72 h, respectively, after OC-2 cells are loaded into the cell culture platform. (b) Cancer cell division is observed in the micro-culture chamber. Note that the image (I) is captured at 24 h after the OC-2 cells are loaded into the cell culture platform and the time interval between the other captured images is 5 min

3.3 High-magnification, real-time imaging of nano-drug-cancer cell interactions and cytotoxicity evaluation

One key feature of this cell micro-culture platform is the ultra-thin culture chamber which allows real-time, cellular, high-resolution imaging at a sub-cellular level by combining bright-field and fluorescent optics. The nano drug is designed by the chemical crosslink of quantum dots with epirubicine by 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC) (Sigma Aldrich, Taiwan) as previously described (Pieper et al. 2000). Briefly, a 100nM sample of 20 nm, 655 nm wavelength emission carboxylated quantum dots (Invitrogen Molecular Probes Q21321MP) suspended in DI H2O mixed with 1.5 nM of Epirubicin, is then added to 1 mM EDC, an amine-carboxyl coupling reagent. After 8-hour incubation, each quantum dot is measured to carry 1,500 epirubicin, on average. The nano drugs complex can be tracked by fluorescent microscopy of the nanoprobes while the therapeutic efficacy can be observed by their morphological change in a time sequence. The toxicity of these QD-epirubicine conjugates investigated using MTT assay reveals that the survival rate of the OC-2 decreases with an increasing concentration of the nanoparticle complex, as shown in Fig. 9(a). The survival rate decreases to 6.9% at a concentration of 10−8 M. Conversely, the OC-2 cells in the medium containing only pure QDs have a survival rate of almost 100%. These results demonstrate that QDs alone do not have an acute toxicity to the OC-2 cells, while the epirubicine conjugated QDs and pure epirubicin have dose dependent cytotoxicity to OC-2 cells. (Figure 9)(b). The LD50 of epirubicin does not show a significant difference after being conjugated with the QDs. All experimental data are repeated five times.
Fig. 9

(a) A comparison of the survival rate of OC-2 cells in the medium containing conjugated QDs-epirubicin and only pure QDs. (b) A comparison of the survival rate of OC-2 cells in the medium containing conjugated QDs-epirubicin and only pure epirubicin

The conjugated QD-epirubicin nano drug is then evaluated for the response of the OC-2 cells upon treatment. The nano-drug complex can be tracked under a fluorescent microscopy while therapeutic efficacy can be observed by morphological change of the cells in real-time using 100X objective oil-lens. Figure 10 shows that the nano-drugs (the red dots) are accumulated around the OC-2 surface initially (Fig. 10(a)), followed by vacuolization of the cytoplasm after 4 h of exposure, suggesting a necrosis type of cell death triggered by this nano drug (Fig. 10(b)). It is noteworthy that nano-drugs surrounding the OC-2 cells are retained even under the continuous flow of the culture medium across the cell culture area (flow rate: 9 μl/hr, driving frequency: 0.12 Hz) during the period of in situ observation. These results suggest that the nano-drugs are strongly attached onto the surface of the OC-2 cells then penetrate into the cytoplasm of the OC-2 cells to induce a necrosis pathway. A wider view by using a 50X objective lens shows that most cells are undergoing necrosis (Fig. 10(c)). These experimental data show that the developed cell culture platform is capable of providing a stable culture micro-environment for long-term, high-resolution, real-time observation at the cellular and the sub-cellular level. Drugs with different concentrations can be injected into the cell culture observation area continuously or in fixed doses by using the integrated micropumps. The interactions between the anti-cancer drugs and cells during a long-term cultivation period can then be investigated with this approach.
Fig. 10

OC-2 Cancer cell response to the administered anti-cancer nano-drug (conjugated QDs-epirubicin) is observed under a microscope. The images (I), (II), (III) are captured by a video camera coupled with a microscope at 0 h, 4 h, and 8 h, respectively, after the nano-drugs are loaded into the cell culture platform. Note that small, red spots in (a) and (b) are QDs carrying anti-cancer drugs. Necrosis of cells after exposure to the nano-drug is observed as vacuolization of cytoplasm can be clearly identified (arrows) (b). The broad view of the 50X objective lens shows that most OC-2 cells are necrotic (c)

4 Conclusions

This study reports a new cell micro-culture platform capable of automatic cell plating, medium supply, waste removal and real-time monitoring of cellular responses to the administered anti-cancer drugs or nanoparticles. The system is comprised of a cell culture module and a micro temperature control module. Continuous medium supply and long-term temperature stability of the cell culture area can be maintained for more than a week. Moreover, the cell culture platform featured an ultra-thin culture chamber for high-resolution cellular and sub-cellular level imaging. Together with these integrated functions, the developed micro-culture platform could become a promising tool for further high-throughput drug screening and biomedical research.


The authors would like to thank the National Science Council in Taiwan for financial support of this project.

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© Springer Science+Business Media, LLC 2009