Contributors: T. Glawdel, C. L. Ren
To date, several microfluidic perfusion culture systems have been developed; however, most rely on external fluid control from off-chip valves and pumps. There are two commonly used pumping methods for microfluidic chips which are pressure driven flow and electroosmotic flow (EOF). Most microfluidic chips for cell culture applications use pressure driven flow as a major pumping method, where in general syringe pumps are used to provide a single source of flow. The flow rate to each cell chamber was fixed by the relative hydrodynamic resistance of the channels. However, pressure driven flow by a syringe pump can suffer from long delay times and a lack of flexibility in multi-function devices. EOF is a compact method of pumping fluids in microfluidic chips where flow rate and direction can be accurately controlled by simply manipulating applied voltages at the reservoirs. However, EOF pumping has several limitations for cell culture: high electric fields, joule heating and high shear stress may be applied to the cells. Consequently, standard EOF pumping which applies an electrical field directly through the working liquid cannot be applied for cell culture applications. These problems can be solved by utilizing the advantages of EOF in electroosmotic pumps (EO pumps) which generates an induced pressure driven flow from EOF flow. EO pumps provide pulse free flow, almost instant flow control, and precise movement of minute volumes of fluid (µL-pL). The integration of internal pumps onto the chip also opens up the possibility of multi-fluid control on the chip without valves. This study investigates the application of on-chip EO pumps to cell culture applications using the pumps to perfuse and sustain cells.
1) The principles of EO pumps.
An electric field is applied in a limited region of the pump generating EOF flow (channel 2). Membranes connect the electrode reservoirs to the fluid flow completing the electric circuit but preventing flow to the electrodes. Due to continuity an internal pressure is generated that pumps the fluid outside the EOF region; a vacuum pressure draws the fluid into channel 1 and a positive pressure drives the fluid through channel 3. The fluid flow in channel 2 is actually a combination of EOF flow and back pressure flow. By placing the cell chamber in the inlet region (channel 1) the electric field will be separated from the cells. As well, the medium will first flow over the cells and through the EO pump eliminating any changes in the medium as it experiences EOF.
2) Designed EO pump for cell culture.
Each EO pump consists of a region of 35 parallel channels (2.78µm H x 100µm W x 1 mm L) that carries the electric field. A series of integrated gel salt bridges provide an electrical contact between the Pt electrodes in the reservoirs and the working fluid in the channels. The gel salt bridge separates the Pt electrodes and the working liquid minimizes electrolysis effects, which may change the pH value of the working liquid and bubbles which may stop current flow.
3) Two EO pumps were integrated into a cell culture chip which can be operated in several configurations.
The first configuration is to run the chip using only the EOF pumps (figure 2.B). In this case, the EO pumps are used to draw the fluid from reservoirs 2 and 3 over the cells in the chamber (flow is from R2->R4, R3->R5). Placement of the pump downstream from the cells avoids the fluid passing first through the pump and then the cells. As well, the electrode near the cell chamber would be GND and the other a negative voltage (V), in this way the cells do not experience a high voltage. In the second configuration (see figure 2.C) flow from a pressure driven source such as a syringe pump provides constant flow over the cells (R1->R4, R1->R5). The EO pumps operating in the same manner as before would divert the flow going to the chambers. For example, if 2uL/min flow is generated by the syringe pump and transported in channel 1, without turning the pumps on the flow will split evenly between channels 2 and 3 (1µL/min). There is little flow through the EO pumps due to the high hydrodynamic resistance of the 35 microchannels. However, if the EO pumps are operating in the same manner as configuration 1 flow will be diverted away from the cell chamber and down the EO pump channels. This provides the possibility of dynamically regulating flow to the cells from a syringe pump without the need for integrated valves. The third configuration (see figure 2.D) is to run the EO pumps in the opposite direction to dispense small amounts of fluid into the stream; for example a discrete amount of steroid. At this point the focus has been limited to configuration 1.
4) Design and operation of the flow network.
The flow network may be described by a 1D circuit model for the flow field. The conditions of using this model include that the flow is laminar and the Navier-Stokes equations become linear, which are true for most planar EO pumps.
5) Fabrication and testing of the chip.
The chips were fabricated in polydimethylsiloxane (PDMS), cell culture tests are currently underway.
Bibliography
T. Glawdel, C. L. Ren, “Electroosmotic Flow Control for Living Cell Analysis in Microfluidic PDMS Chips (PDF)”, Mechanics Research Communications: Special Issue of Recent Advances in Microfluidics, 36 (2009) 75-81.