As PDMS and glass are electrical insulators, it is a reasonable expectation that an uninterrupted microchannel filled with a salty buffer would drastically increase its electrical resistance when it is pinched off by a closed valve, since the insulating material of the valve would interrupt the electrical circuit of the channel. Simple preliminary experiments confirmed that prediction.
The observed effect establishes a 1-1 correspondence between channel resistance (low/high) and valve status (open/closed). The mechanical properties of the valve establish another 1-1 correspondence, between the valve status (open/closed) and the applied pressure (insufficient/sufficient). The status of a single valve does not report the value of the applied pressure, but reports if the applied pressure is below or above the valve’s characteristic pressure. In essence, a single valve status produces a single inequality.
Valves of different dimensions have different characteristic pressures.14
Therefore a heterogeneous array of valves, acted upon by the same applied pressure, would report a set of inequalities, whose bounds are the different characteristic pressures. Thus the magnitude of the applied pressure would be limited to one of a set of intervals delineated by these bounds.
This analysis suggested the basic scheme described above. The scheme was implemented for applied pressures of 0.5 to 19 psi above atmospheric, in steps of 0.5 psi. The maximal value was set by the pressure at which all valves were closed. Further increase in the pressure would not change the result, while pressures in excess of 20 psi significantly increase the risk of layer delamination. Each flow-channel prong had low resistance (26–64 MΩ) when its valve was open, while the resistance increased beyond the dynamic range of the multimeter (2 GΩ) when the respective valve closed completely. shows the resistance pattern versus applied pressure. Since in certain cases multiple valves closed at the same pressure, their resistance patterns overlapped and are represented by the same color.
FIG. 2 Device Function. Each valve closes when the applied pressure exceeds the valve’s characteristic pressure. Thus the status of each valve determines applied pressure to an upper or lower bound. Then the status of a heterogeneous set of valves produces (more ...)
This system can be used as a pressure gauge. For example, if all valves up to and including “blue” in are in “high resistance” state, while all valves above it (light green, red, black) are in “low resistance state,” then the applied pressure is between 10.5 and 11 psi.
The precision of the measurement is dependent on the spacing of curves in and thus, on the spacing of characteristic pressures. In the demonstrated device, valve width, thickness, and material were kept constant, while valve length was varied between 55 and 200 µm. However, if that restriction is relaxed, many more values for characteristic pressure become available.14
Then the pressure spacing can be shrunk accordingly, thereby improving the precision of the measurement.
Meticulous mappings of the phase space of closing pressures versus valve dimensions14
showed that individual valve behavior is robust, reliable, and reproducible, thereby attesting to the robustness, reliability, and reproducibility of its derivative devices, such as the presented pressure gauge. Hence, the accuracy of the measurement is set by the accuracy of the digital gauge and the size of the pressure step used in calibration.
The response time of the overall system is governed by the mechanical actuation time of the valves and the time taken by the electrical measurements. Micropump characterization24
sets the former at the scale of milliseconds, while the latter typically are orders of magnitude faster. Thus the overall sensor response time would be milliseconds. Such speeds are adequate both for isolated measurements and for continual monitoring in pressure regulation loops.
In the presented embodiment, the pressure gauge achieves large signal-to-noise in valve status measurements, through the use of a salty buffer. Hence, there are concerns that salt ions could diffuse through the elastomer, thereby adversely changing the ambient conditions in particular biochemical applications. However, biochemistry typically works with salinities at physiological level or above, which is essentially what we used here. Furthermore, in any particular case, the electrical measurement fluid can be set at the same salinity as the biochemical environment to prevent nonzero net ionic flux. Finally, potential losses of signal to noise in valve status measurements can be offset by optimizations of the geometry of the conducting channels.
Other pressure-sensing schemes and devices have been proposed and demonstrated. A microfluidic differential manometer has been used to study cells.15
However, it does not produce an absolute measurement of the pressure, while the readout is done visually and thus cannot be straightforwardly automated or miniaturized.
sensors have convenient electrical readout and are miniaturizable. However, it may prove challenging to adapt these devices for inexpensive mass production, robust performance, and integration into particular systems, especially since many biological and biomedical applications already place tight constraints on the properties of the substrate, e.g., due to binding surface chemistries5
and/or cell toxicity. In addition, some of the devices18,20
have inadequate dynamic ranges for typical elastomer microfluidics.
Vacuum-sealed silicon sensors21
utilize capacitance, piezoresistivity, or resonance measurements involving a silicon microdiaphragm. These sensors are compatible with mass production, are electrically readout, and are at the correct physical scale. However, the capacitance version requires voltages that are impractical for overall system miniaturization, the piezoresistive version is limited to silicon substrates, and the resonator versions are difficult to fabricate and are also limited to silicon substrates.
A pressure sensor,22
involving an interdigitated capacitor with metal oxide dielectrics, has high sensitivity but excessive physical size and insufficient dynamic range. An optical sensor23
measures the deflection of a silicon membrane by the intensity of light reflected from it; however, a light source, waveguides, and a photodetector are necessary, and so, overall system miniaturization is problematic.
By contrast, our device is simple, inexpensive, easy to fabricate, and straightforwardly integrable within elastomer chips. In principle, it is compatible with any substrate, since buffer-filled microchannels and vias6
can be used as three-dimensional electrical connections to access the array. In addition, the system is reliable, robust, reproducible, and appropriately sized. It requires simple electric circuitry to function and offers adequate dynamic range, accuracy, and precision. So far, it seems to be the best overall solution for a pressure sensor integrated in elastomer microfluidics. The envisioned particular application is to monitor pressure and electrically report the result to logic circuitry that controls the overall PDMS MEMS. This capability would be essential in completing the pressure-control loop when pressure generation (for valve actuation and fluid transport) is achieved within the integrated chip of the future.