Cell-Based Microfluidic Biochip for the Electrochemical ... - CiteSeerX

Reference of Transducers Conference: AO0165

Cell-Based Microfluidic Biochip for the Electrochemical Real-Time Monitoring of Glucose and Oxygen N. Pereira Rodrigues a*, Y. Sakai a,c, T. Fujii a,b a b c

LIMMS CNRS/IIS, University of Tokyo, Tokyo, JAPAN

Institute of Industrial Science, University of Tokyo, Tokyo, JAPAN

Center for Disease Biology and Integrative Medicine, University of Tokyo, Tokyo, JAPAN

ABSTRACT We report the fabrication of a microfluidic biochip dedicated to realize a precise and realtime monitoring of the composition and modulation of the cell medium by comparing the rates of glucose before and after contact with the cells. The device was composed of specific glucose and oxygen amperometric sensors integrated at the inlet and outlet microchannels of a poly(dimethylsiloxane) (PDMS) cell-chamber. The microfluidic device was designed to be compatible with cell cultures and the performances of the integrated sensors in the dynamic conditions of liquid flow were reported during calibration. We show here the compatibility of the developed biochip with cell culture analysis in the case of hepatocytes (HepG2) and put in evidence the good performances of the integrated microsensors in terms of linearity of response and sensitivity of detection during calibration.

Keywords: integrated amperometric microsensors, PDMS microfluidic biochip, glucose, oxygen.

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1.

Introduction

In recent years, there has been intensive research in the field of medical diagnostics aimed at the fabrication of reliable analytical systems for multiple and accurate analysis of in-vitro steady-state and transient fluxes of molecules involved in cellular metabolism [1,2]. In the field of cell analysis, there is an increased demand for microdevices dedicated to the monitoring of specific analytes involved in cell growth and development, which can thus contribute to a better understanding of the factors that influence metabolic processes implied in several diseases. This requires the fabrication of devices compatible with cell culture by microscale techniques. These cell-based biochips are expected to perform in-vitro analysis by mimicking in-vivo conditions. Microfluidic devices obtained by photolithographic and planar techniques are now widely used because of their numerous advantages such as sample holding, reagent mixing, separation and detection [1,2]. Besides, these devices are adapted to cell culture analysis because of their potential to modulate the biochemical composition of the medium surrounding the cell culture. Electrochemical miniaturized sensors are particularly attractive for various in-vivo and ex-vivo biomedical applications, including continuous in vivo monitoring, measurement of analytes in extremely small volumes, monitoring of localized events or biosensing resistive media [3]. We report the fabrication of a miniaturized cell-based biochip dedicated to perform toxicological analysis on cells during their growth and development. For that purpose, electrochemical sensors are integrated at the inlet and outlet microchannels of a PDMS cell chamber to perform real-time and continuous monitoring of glucose and oxygen during cell culture. Previous works put in evidence the electrochemical monitoring of these species in the flow dynamic conditions by using external FIA (flow injection analysis systems) [4-6].

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However, biochips integrating electrochemical sensors dedicated to cell activity analysis have not been reported so far. In this study, the design of this biochip is detailed and the biocompatibility of the all biochip is discussed in the case of HepG2 cell culture. Calibration of the integrated sensors in biocompatible conditions of flow dynamics was performed without any cells in the device and the performances of the integrated sensors (sensitivity, stability and linearity of response) are discussed. To validate the originality of biochip design, simultaneous detection of glucose or oxygen sensing at the inlet and outlet microchannels was performed.

2.

Materials and Methods

2.1. Design of the Biochip The developed biochip shown in Fig.1 is composed of two arrays of glucose electrochemical sensors on a glass substrate and of a PDMS microfluidic device. The biochip is designed to perform continuous real-time monitoring of the composition of the cell culture medium by comparing the rates of dissolved oxygen and glucose before and after contact of the flowing medium with the cells. For that purpose, oxygen and glucose amperometric sensors were integrated at the inlet and outlet of the PDMS cell chamber. The dimensions of the cell-chamber (1cm2 area, 300μm height) were chosen to allow the culture of ~2-4×105 inoculated human hepatoblastoma cells (HepG2). The final dimensions of both inlet and outlet microchannels were 800 µm large and 300 µm high. Therefore, the flow dynamics inside each channel was designed to present similar flow velocities at the position of each array of sensors. The connections between the microdevice and the external fluidic system were done with Teflon and silicon tubes (Upchurch Scientific). The perfusion system was composed of a peristaltic micropump (Maxon DC motor) and a 6-entries micro-splitter valve (Upchurch Scientific).

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2.2. Thin-film electrodes fabrication First, a glass substrate (Matsunami, 1.8mm thick) was cleaned during 15min. After rinsing and drying, 50nm Ti and 200nm Au were sequentially deposited by evaporation. The electrodes geometry was then defined by a conventional photolithography process by spincoating a 35µm thick insulating layer made of S1813 positive photoresist (AZ Photoresist) with 3000rpm for 30s. The reference electrode was fabricated by evaporating 200µm Ag. Another spin-coat of 5µm of AZP1150-90 (Zeon Corporation) negative photoresist was performed and the final dimensions of the Ag electrodes were obtained by wet etching. An AgCl layer was obtained by dipping the thin-film Ag electrode in a FeCl3 solution. The biochip was composed of two arrays of thin-film microelectrodes (at the inlet and outlet of the cell chamber). Each array of electrodes was split up into two groups of three electrodes (one group dedicated to oxygen and the other one to glucose monitoring) and was composed of one gold-modified working electrode, one Ag/AgCl reference electrode and one gold counter electrode. The active area of the working electrodes was 0.16mm2 and 0.016mm2 in the case of glucose and oxygen respectively.

2.3. Description of the integrated glucose and oxygen sensors The surfaces of the working electrodes were modified by specific membranes to achieve a selective and sensitive amperometric detection of glucose and oxygen: -

The glucose sensors were based on the immobilization of GOx (from Aspergillus Niger type X-S, 100000-250000 U/g, Sigma) on the surface of the Au working electrodes. Nafion® (5% alcoholic solution, Aldrich) is a biocompatible polymer has been widely used in biosensors fabrication as a protective and selective coating material, as well as a support for enzyme immobilization [7-9], and was therefore

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used as a matrix for GOx. A 2µl-mixture of 5g.l-1 Nafion® and 0.2g.l-1 GOx (38U.ml1

) was spread onto the Au working microelectrodes and was allowed to evaporate at

room temperature during 30min. -

Oxygen sensors were achieved by modifying the surface of the electrodes with Nafion®, which is in this case a permeable membrane that also protects the electrodes from fouling, enhances their lifetime and induces a better sensitivity of oxygen detection [10]. 1.5µl of diluted Nafion® alcoholic solution (containing 20% H2O) were deposited on the surface of the working and counter electrodes dedicated to oxygen monitoring.

Sensors calibrations were with a flow rate of 30μl.min-1 at 25oC without the cells inside the biochip.

2.4. PDMS microfluidic device fabrication PDMS (Dow Corning) is a biocompatible material, with interesting properties such as optical transparency, rapid prototyping and good sealing properties to glass [11,12]. The PDMS microfluidic structure was fabricated through replica molding processes with a master that contained the negative pattern of the structure. After the cleaning step, a negative master was fabricated with a SU8-2075 photoresist, spin-coated on a Si wafer by using a conventional photolithography process [13]. The height of both the channels and cell chamber was 300μm. CHF3 plasma (RIE-10NR, Samco) was applied to deposit a fluorocarbon layer onto the obtained SU8-mold for future easy release of the PDMS layers [14]. Then, the liquid-state PDMS was poured after removing all bubbles by putting the master and PDMS in a vacuum chamber. The solidification step consisted of baking the PDMS at 70°C during 1h30. The PDMS was then peeled off from the master and the inlet and outlet of each PDMS layer were drilled with 1mm diameter punch. After modifying the

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surface of the working electrodes, as previously described, the glass substrate containing the array of electrodes and the PDMS chamber were aligned and permanently stacked by O2 plasma (RIE) treatment.

2.5. Measurement procedure All chronoamperometric measurements were performed with a four channel multimode potentiostat (ESA Biosciences Inc.) linked to a PC-based data acquisition (Biostat). No special electrochemical pre-treatment of the electrodes was performed before measurement. All measurements were collected at room temperature at E = 0.65 V vs. Ag/AgCl for glucose and E = -0.60 V vs. Ag/AgCl for O2. A physiological Dulbecco’s Phosphate Buffer Solution (pH = 7.4) was used as a perfusion fluid for calibration.

2.6. Cells and culture medium Human hepatocarcinoma liver cells (HepG2), obtained from the Japanese Collection of Research Bioresources (JCRB), were inoculated in the cell chamber of the biochip. The culture medium was Dulbecco’s Modified Minimum Essential Medium (DMEM, Nissui Pharm. Co. Ldt.) supplemented with 10% fetal bovine serum, 2% HEPES, 1% antibiotic/antimitotic, 1% MEM-NEEA and 1% ascorbic acid.

3.

Results and Discussion

3.1. Cell culture in the biochip Because the present biochip is designed to be used for cell analysis, it is paramount that the validation of the device should be done in conditions required for cell culture. Therefore, one main important issue to be addressed is the flow dynamics in the device. Thus, the metabolic cell functions and morphology are particularly sensitive to the shear stress induced

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by the flow dynamics of the microfluidic environment [15-17]. In particular, the perfusion flow velocity should be kept under a certain value in order to avoid the detachment and/or the accelerated death of the cells inside the device. Tanaka et al [16] reported a maximum value of shear stress of 6dynes.cm-2, under which no detrimental effect on hepatocyte cell cultures occurred. In the present work, we decided to set and maintain the flow dynamic conditions close to those used by Leclerc et al. who studied HepG2 (Human hepatocellular liver carcinoma) cell lines with a shear stress in the cell-chamber of ~0.03dynes.cm-2 [17]. The wall shear stress (τω) is directly related to the velocity of the flowing medium according to the modified Poiseuille-Hagen equation:

τω = (μ x 4u) / R

(1),

where μ is the medium viscosity (Pa.s), u is flow velocity (m.s-1) and R is the channel diameter (m). τω is expressed in Pa or dynes.cm-2 (1Pa=10dynes.cm-2). For that purpose, we performed simulation of the fluid flow dynamics inside the device (COMSOL Femlab 3.3). We used as a mathematical model the steady-state incompressible Navier-Strokes equations. The flow velocity inside the cell-chamber, uch, .that corresponds to a shear stress of τω~0.03dynes.cm-2, is uch=2.5×10-4m.s-1. We found that the flow velocity to be set at the entrance of the inlet microchannel, uin, for a maximum value of uch~2.5×10-4m.s1

inside the cell-chamber, is uin~3.5×10-3 m.s-1. This velocity corresponds to a total flow rate

of ~30μl.min-1. Therefore, in the following, all experiments are performed with a flow rate of 30μl.min-1. A photograph of HepG2 cells directly inoculated inside the cell chamber of the biochip (initial concentration of cells: ~2x105 cells.cm-2) is shown in Fig.2. After 5 days, we observed that the cells exhibited a good development and activity due to their 3-dimensional growth and hexagonal shape. Therefore, the configuration of the PDMS device is totally in

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accordance with HepG2 cell culture in terms of biocompatibility and shear stress considerations.

In the following, we validate the performances of the glucose and O2 integrated sensors by using these dynamic flow parameters without having the cells inside the cell chamber.

3.2. Glucose monitoring: performances of Au/Gox-Nafion® The strongest requirement for the present device to be well-adapted for future studies of cell culture is to ensure that the sensitivity of the Au/GOx-Nafion® biosensors is in agreement with the typical ranges of glucose consumption found in literature for biological monitoring of glucose during cell culture [6,11,17,18]. Leclerc et al [17] found that the daily glucose consumption of HepG2 cell lines was of the order of 1.5mg.day-1, i.e. 5.8×10-9 mol.min-1, for 104 inoculated cells per cm2. Given the previously value of flow rate (30µL.min-1) in the microchannels and the inoculated cell concentration (2-4×105 cells per cm2), the molar concentration of the glucose solutions to be tested is in the range of 1.4-2.8 mmol.l-1.min-1. Though it is difficult to estimate precisely the glucose consumption by the HepG2 cells to be integrated later, we considered the above values for the validation of the performances of the Au/GOx-Nafion® glucose sensors. To evaluate the sensitivity of the Au/GOx-Nafion® sensors in the flow dynamic conditions, experiments were first performed by measuring the glucose concentration at the inlet channel. After polarization of the electrodes in PBS flow, a stable current was initially reached after 15 min, and the current was then quickly stable when the electrode was polarized again. Fig.3 shows the on-line response of the inlet Au/GOx-Nafion® sensor by successively injecting 80μl of different concentrations of glucose (from 10 to 2mmol.L-1) in the fluidic system. Between two glucose injections, only PBS was flowing. We observed an increase of

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the intensity of the anodic current proportional to the concentration of glucose injected in the perfusion system (inset of Fig.3). Moreover, all subsequent injections of identical concentration lead to both identical peak intensity response and recovery to the initial intensity, with identical response and recovery times. Therefore, this result shows that Au/Gox-Nafion® integrated sensors yield in the current experimental conditions a great linearity of response with a sensitivity of 0.5 ± 0.1 nA.mmol.L-1, with a detection limit of 120µmol.l-1. The integrated Au/GOx-Nafion® sensors yield sufficient sensitivity for glucose measurement and therefore meet perfectly the requirements imposed by the expected cell culture glucose consumption of hepatocytes. We point out that possible differences of detection of sensitivity between the inlet and outlet sensors are of minor importance. Indeed, in future works with cells, the monitoring of glucose will be directly expressed in glucose concentration after calibrating each Au/GOx-Nafion® sensor. Because the present device is designed for cell culture studies, which require glucose monitoring during several hours, the stability in time of the glucose sensing is a parameter of great importance. For that purpose, we tested the sensor for several hours with a continuous flow of constant concentration. The experiment showed in Fig.4 was done by monitoring simultaneously the inlet and outlet rates of glucose in the case of transient and continuous injections of 10mmol.l-1 of glucose. We observe that both inlet and outlet Au/GOx-Nafion® glucose sensors could reach a steady-state current reached after 2 min and exhibit a similar response and stability. Besides, we tested the sensors in PBS at 37oC and found that the response of Au/Gox-Nafion® exhibit a similar range of sensitivity (data not shown). Therefore, these experiments show that the integrated Au/GOx-Nafion® electrodes are well adapted to the monitoring of glucose in the dynamic flow conditions and yield great stability. The signal stability indicates in particular that no significant variation of the GOx activity occurs over time. It is noteworthy to point out also that sensors could be used over

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several days, by stocking the biochip at 4oC, in order to stabilize GOx activity. We point out that, as in the case of glucose monitoring, possible differences of detection of sensitivity between the inlet and outlet sensors are of minor importance. Indeed, in future works with cells, the monitoring of glucose will be directly expressed in glucose concentration after calibrating each Au/GOx-Nafion® sensor.

3.3. Oxygen monitoring : Performances of Au/Nafion® As for glucose monitoring, the strongest requirement for the present device to be welladapted for future studies of cell culture is to ensure that the sensitivity of the Au/Nafion® sensors is in agreement with the expected range of oxygen concentration. The average oxygen uptake rate for HepG2 cells is in the range of 50-237μmol.l-1 (ie. 4-21%) [19]. To calibrate and evaluate the linearity of response of Au/Nafion® oxygen sensors, dissolved O2 concentration in the flowing medium was modulated by bubbling the flowing solution with pure N2 and O2 [20]. The exact amount of dissolved oxygen was determined by connecting commercial O2 meters (Strathkelvin Instruments, UK) at the inlet and outlet of the device. Fig.5 shows the calibration curve obtained for the inlet Au/Nafion® oxygen sensor by modulating the concentration of O2 from 50 to 500μmol.l-1. The results shown Fig.5 put in evidence a perfect concordance between the amperometric signal of Au/Nafion® sensors and the dissolved O2 concentration determined by the O2 commercial meter connected to the flowing circuit. This experiment was performed several times in a long time period, and no degradation of the electrodes was observed in the presence of Nafion® (data not shown). Besides, these results show that there is a perfect concordance between the responses of the integrated sensors and the oxygen meters, which leads to an excellent linearity of response of the integrated Au/Nafion® oxygen sensor in a larger range of O2 concentration (inset of Fig.5). The sensitivity of the sensors was found to

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be of 0.5± 0.05nA.mmolO2.L-1 in the microfluidic conditions without any degradation. As for glucose, O2 monitoring could be performed during hours without degradation of the electrodes, and the sensor could be used for several days. In future works with cells, the monitoring of oxygen will be directly expressed in dissolved O2 concentration after calibrating each Au/ Nafion® sensor. In way to validate the configuration of cell culture studies, we performed the simultaneous monitoring of inlet and outlet rates of oxygen by modulating O2 concentration from 237μmol.l-1 to 50μmol.l-1. The results obtained Fig.6 show an excellent concordance of the sensors, exhibiting the same sensitivity (0.50±0.05nA/μmol.l-1). Therefore, these experiments show that the integrated Au/ Nafion® electrodes are well adapted to the monitoring of dissolved O2 in the dynamic flow conditions and yield great stability and short time of response. We also point out that the flow rate of the device is of great importance for O2 sensing. Indeed, due to the high permeation of O2 through the PDMS, we needed to find a compromise between the flow velocity in the device and the kinetics of permeation of PDMS while modulating the composition of O2 in the flowing medium. Therefore, we can conclude that in this configuration of device (dimensions of the channels and cell chamber, position of the sensors), and in these experimental conditions (total flow rate), comparison of O2 concentration between inlet and outlet was successfully realized. Complementary experiments at 37oC are now in progress.

4.

Conclusion and Outlook

In summary, we reported the performances of a biocompatible PDMS microfluidic biochip dedicated to cell culture analysis for the real-time electrochemical monitoring of glucose and oxygen. The device was designed and tested in conditions to be compatible with cell cultures. The overall performances of the device for glucose and oxygen monitoring

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were tested and we showed that the integrated Au/GOx-Nafion® and Au/Nafion® electrodes exhibited accurate, sensitive, stable, and reproducible response in the given dynamic flow conditions. We performed the simultaneous inlet and outlet monitoring of oxygen and glucose, and showed that the developed biochip is adapted to the study of transient effluxes of these species during cell culture by comparing the rates of these species before and after contact with the cell chamber. Therefore, the present biochip opens promising prospects for toxicological studies on cells concerning the cellular metabolic activity via O2 and/or glucose sensing, but also the cellular system evolution following different environmental compositions. Studies implying detection in cell culture medium with HepG2 cells are at the moment in progress.

5.

Acknowledgements

We would like to acknowledge CNRS, JSPS and MEXT for their financial support.

6. [1]

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Legends:

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Fig.1. (a) Cross-section and (b) general schematic view of the developed biochip composed of two arrays of glucose and oxygen electrochemical microsensors integrated at the inlet and outlet microchannels of a PDMS microfluidic chamber.

Fig.2. Photograph of HepG2 cells in the PDMS cell chamber of the biochip (5 days after inoculation of 2x105 cells)

Fig.3. Real-time amperometric response of the inlet Au/Gox-Nafion® glucose sensor by adding defined glucose concentrations in the biochip (pH=7.4, Flow rate: 30μl.min-1). Inset: Au/Gox-Nafion® calibration curve (I=f([Glu]).

Fig.4. Simultaneous response of the inlet and outlet Au/GOx-Nafion® glucose sensors after successive additions of (a) 60μl of 10mmol.l-1, (b) only 10mmol.l-1 of glucose, and (c) only PBS. The dashed lines correspond to the injection times.

Fig.5. Comparison of Au/Nafion® O2 sensor response (black line) with a commercial meter connected to the perfusion system (gray dotted line) when modulating the flowing medium composition with pure O2. Flow rate: 30μl.min-1

Fig.6. Simultaneous response of the inlet and outlet Au/Nafion® O2 sensors when modulating the flowing medium composition with pure N2. Flow rate: 30μl.min-1.

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