Pavel
Neuzil
*,
Juergen
Pipper
and
Tseng Ming
Hsieh
Institute of Bioengineering and Nanotechnology, 31 Biopolis Way, Nanos #04-01, Singapore 138669. E-mail: pneuzil@ibn.a-star.edu.sg
First published on 17th May 2006
We have designed, fabricated and tested a real-time micro polymerase chain reaction (microPCR) system. It consists of a microscope glass cover slip placed on top of a micromachined silicon chip integrated with a heater and a temperature sensor. A single μL of a sample containing DNA was placed on the glass and encapsulated with mineral oil to prevent the evaporation of water, thus forming a virtual reaction chamber (VRC). The PCR chip required half a second to heat up from 72 to 94 °C and two seconds to cool from 94 to 55 °C, corresponding to a cooling rate of −20 K s−1. The real-time PCR yield was determined by a fluorescence method. The melting curve analysis method as well as capillary electrophoresis was performed to determine the purity of the PCR product. As the glass slip is disposable, cross-contamination from sample to sample is eliminated. The total cost of running the PCR is given by the value of the cover slip and its treatment.
It is important that the PCR system3 has good temperature control, maintains temperature uniformity within the sample and has a sample heating (cooling) rate of at least 5 K s−1 (−5 K s−1). Sample to sample cross-contamination should be also avoided. Temperature control is typically performed by a feedback loop system, while temperature uniformity is achieved by highly thermally conductive but bulky materials such as copper. A high heating rate is accomplished by the implementation of a proportional integrated derivative (PID) control method limited by maximum dissipated power and heat capacitance. A high cooling rate is rather difficult to achieve and bulky systems require forced cooling by either a thermoelectric element4 (often called a Peltier element) or by other means, such as water.5 They end up being complicated and even more power hungry devices than conventional PCR systems.
As the systems are bulky, their thermal time constants are in minutes rather than seconds. That results in long transition times and unwanted by-products of the PCR. The high power consumption eliminates the possibility of making a battery-operated and portable PCR system. In addition, the reaction tubes are large and the required amount of PCR cocktail makes the whole process expensive. Also, the detection of PCR product has to be done off-line, i.e. in another instrument, resulting in additional cost.
µPCRs systems can be also categorized based on the heating system, which is either direct or indirect. Direct heating PCRs have the heater as well as the temperature sensor integrated with the device. Indirect heating can be performed by infrared radiation.13 PCR chips with integrated heaters/sensors are more complicated that ones with indirect heating but overall, the whole system is more compact and better-suited for small, portable PCR systems.
Currently, the three most popular materials for μPCR fabrication are silicon, glass and plastic. Silicon is an excellent material for a thermal-cycler chamber. It has a high thermal conductivity λ of 157 W m−1 K−1 and once it is thermally isolated, the chamber has good thermal uniformity.10 Also, micromachining is well established for this material. The drawback is that the silicon surface itself inhibits the PCR and its surface has to be covered with another material, such as silicon oxide or silicon oxynitride.
There have been attempts to make PCR chips on a glass substrate with a thin14 or thick film metal heater and sensor. Glass has a thermal conductivity λ of 1.1 W m−1 K−1, more than a hundred times lower than that of silicon. Due to its low thermal conductance, the systems made of glass are thermally isolated. However, creating a microfluidic system by glass machining is rather difficult compared to either silicon or plastic processing.
The obvious solution is to use micromachined silicon capped with another chip, which can be made of either silicon or glass.15 A short response time and low power consumption was achieved with systems made of both materials. The drawback is that the micromachining technique is relatively complicated and consists of 4–6 levels of lithography, including silicon etching and wafer to wafer bonding. Nevertheless, the combination of glass and micromachined silicon gives the designer two materials with a large difference in their thermal conductance, which allows tailoring of thermal and mechanical properties.
The third material commonly used for PCR is plastic (such as polycarbonate),16 polydimethylsiloxane (PDMS)17 or a composite used for printed circuit boards (PCB).18 All those materials have a cost advantage over both silicon as well as glass and they are simple to process. Polycarbonate can be shaped by a hot embossing technique,16 while PDMS polymerises in a mold.17 PCB technology is also well established. The common drawback is the low thermal conductivity of the plastic, and thus it should be combined with either metal or silicon in order to achieve the desired thermal properties.
A simplified approach was recently19 proposed, where the reaction chamber was made by encapsulation of a water based sample in oil, forming a virtual reaction chamber (VRC). As no solid cover or microchannels were required, the device fabrication then consisted only of deposition and patterning thin film heaters and temperature sensors on a suitable substrate, which is still too costly for a disposable system.
In order to eliminate cross-contamination between samples, the safest way is by using a disposable system. At the very least, the part of the device, which comes into contact with the sample should be disposable. So far, many different systems have been proposed. These systems typically do not fulfill all the requirements listed above and they are relatively expensive. An approach with a disposable part made of a plastic sheet was presented earlier.20 A set of wells was formed by hot embossing and the whole set was placed on top of the heaters. This system employs a relatively complicated microfabrication process and the disposable plate has to be customized. There is a need for a μPCR that is simple to manufacture, easy to operate, and economical enough to be disposable. The optional ability to be integrated into a complete μTAS system is highly desirable.
Here we report on a system (see Fig. 1), which fulfills all those demanding requirements. We propose to place a VRC14 on a microscope cover slip sitting on a micro-machined silicon structure integrated with both a thin film heater and a sensor. For the integrated sensors, the resistance temperature detector (RTD) type of sensor is a natural choice due to the simplicity of its fabrication, as there is only one metal deposition and two lithography steps required. Furthermore, the disposable cover slip part is not subject to any processing. The PCR system still maintains a high heating rate of greater than 30 K s−1 and a cooling rate of −20 K s−1.
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Fig. 1 A photograph of a PCR chip soldered to a printed circuit board (PCB) with a square microscope glass cover slip. Four water-based samples, each with a volume of 1 μL, are placed on the glass and covered with 5 μL of mineral oil. The samples contain blue dye for easier visualization. Due to the soldering technique, we have achieved low resistance between the PCB and the chip as well as robust mechanical connection. Also, handling the whole PCB is much simpler than dealing with bare silicon chips. |
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Fig. 2 Schematic cross-section of an individual heater (top) with a VRC, which is water-based droplet enclosed by oil, placed on the top of the heater. Both the heater as well as the sensor are made of a thin film of metal and thermally connected by placing them at the same silicon ring. Silicon is separated from the VRC by a thin sheet of glass, which transfers the heat from the heater to the sample. Schematic drawing of a second version of a thermal cycler (bottom). The lighter part is micromachined silicon with a thickness of 450 μm while the darker donut shapes represent the heater (outer) and sensor (inner). The temperature distribution is symmetrical along the axis of symmetry represented by the dashed line. Thermal conductance is given by the beam material, its length and cross section. Thermal capacitance is given by the double donut volume together with the volume of the VRC. |
Due to the high thermal conductivity of silicon compared to that of glass, the thermal cross talk between chambers is expected to be minimal. This would allow running of a number of PCR systems independently in close proximity to each other.
The silicon device is symmetrical along the cantilever axis. We decided to take advantage of the symmetry and create a double donut shape (see Fig. 2) with the heater and sensor located at the inner donut. Both donuts are connected by two beams with thermal conductivities of at least half that of the inner donut. It results in temperature uniformity within the inner donut. The heat was dissipated almost entirely via the cantilevers, supporting our expectation that the cross talk between zones will be minimal. This design allows controlling the temperature of all four areas independently from each other, and thus we could run four different PCR protocols simultaneously.
The chip was designed with a bonding pad configuration identical to a standard leadless chip carrier (LCC-68) socket, so that it could be clamped into a conventional testing socket to determine device thermal parameters. Since our device thickness was only 0.45 mm as compared to the standard LCC chip, which has a thickness of about 3 mm, we added a plastic frame on the top of our chip to compensate for its smaller thickness.
The resistance value R of a resistor versus temperature difference ΔT can be described by a simplified equation:
R = R0(1 + αΔT) | (1) |
Sensor resistance | 320 Ω |
Heater resistance | 110 Ω |
Sensor TCR | 0.33% K−1 |
Sensor sheet resistance | 0.11Ω/□ |
Unit heat conductance | 4.40 mW K−1 |
Unit heat capacitance | 6.60 mJ K−1 |
PCR thermal time constant | 1.74 s |
The thermal behavior of any system is described by the differential heat balance equation:
![]() | (2) |
Previously, a pulse method to derive thermal parameters of bolometers for infrared detection was published.23 In principle, bolometers exhibit a behavior similar to PCR devices, so we have used an identical testing method. The sensor under evaluation together with three external resistors formed a balanced Wheatstone bridge. It was powered by pulses with durations of 1 ms and a voltage amplitude of 5 V with a repetition rate of 1 pulse per second. A DC voltage signal with an amplitude between 0 and 1 V was superposed onto the pulses.23 The thermal capacitance H of the device was calculated from the derivative of temperature with respect to time. The temperature increase above ambient due to the applied DC voltage is a function of thermal conductance G. Obtained values of H and G were verified by direct measurement of the system’s time constant τ (equal to H/G). All measured and calculated electrical and thermal parameters are listed in Table 1.
The PCR device was connected to the temperature control electronics as described later in this paper. The temperatures of the individual heaters were set approximately to 65, 85 and 94 °C and infrared (IR) images at wavelengths from 8 to 12 μm were captured. The camera’s temperature resolution was 0.1 K of the noise equivalent temperature difference (NETD) (see Fig. 3). As shown in Fig. 3, the temperature variation across the heaters is less than 1 °C, and thus the device is well-suited to perform the PCR sequence.
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Fig. 3 An infrared image of a PCR device soldered to the PCB is shown here. Blue represents the cold zone and red represents the hot zone. The scale is shown on the right hand side. The bottom heater temperature was set to 95 °C, left and top to 85 °C and right heater temperature to 65 °C. All heaters are covered with a glass slip with a diameter of 12 mm. The temperature profile along the green line was extracted. The temperature variation within both zones of interest is within ±0.5 °C. It verifies the assumption as well as the FEA simulation results that the thermal non-uniformity within the zones matches the requirements for the PCR system. |
The temperature sensor, together with two fixed and one adjustable resistor, formed a Wheatstone bridge. Its outputs were connected to an INA143US (Burr-Brown, Inc.) differential amplifier with a fixed gain of 10. Its output was linked with LabView software by the same card as the one which controlled the IR 2121. The complete PCB consists of four individual channels to run four PCRs in parallel.
The PCR device was calibrated to a temperature precision of better than 0.5 °C. The device calibration was performed in a temperature-controlled bath filled with Fluorinert™ 77 (3M Specialty Chemicals Division, Inc.). Its temperature was measured by temperature sensors TSic™ (IST-AG, Switzerland) calibrated with a precision of 0.1 °C in the range from 50 to 100 °C, soldered at the PCB next to the PCR device.
The output values from all four channels were stored in a LabVIEW setup file and used for the feedback measurement. The microscope glass cover slip was placed on the PCR chip. A VRC sample with a volume of 1 μL and 5 μL of oil was dispensed above the heaters (see Fig. 1). The above procedure precisely verified the temperature of the heater but not of the PCR sample itself, which could be different. The sample temperature was determined by melting curve analysis,24 which is described later in this paper. We found that the sample temperature is two degrees lower than the temperature of the heater at 94 °C and the setup file was corrected accordingly.
We are currently developing a fast PCR system capable of 40-cycle process in 5 min or less,27 which is not subject of this paper.
For melting curve analysis24 the sample was cooled down to 65 °C for 1 min after which the temperature was continuously raised to 95 °C with a heating rate of 0.01 °C s−1. During the operation, both the fluorescence signal as well as the temperature sensor value was recorded simultaneously.
The next step was the calculation of an average value of the fluorescence signal during the end of the extension phase at 72 °C. In order to extract the fluorescence output signal from the PCR cycles, a short program was prepared using Fortran. The program input parameters were the center of the first data block shown by a blue arrow in Fig. 4, the length of the data interval and the number of intervals. The program then averaged the signal from the interval and associated it with a cycle number for all 50 cycles. The same procedure was repeated for the fluorescence signal at 94 °C (indicated by the red arrow in Fig. 4) in order to obtain a baseline signal to be subtracted from the PCR output signal at 72 °C. The subtracted data set was approximated with a sigmoidal function:
![]() | (3) |
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Fig. 4 Fluorescence signal detected from a PCR chip versus time over 50 cycles. The fluorescence signal from a single PCR cycle is shown in the inset. The real time PCR data points were obtained by subtracting the fluorescence signal at a temperature of 94 °C (shown by a red arrow in the inset) from the signal at a temperature of 72 °C (shown by a blue arrow in the inset). The cycle threshold (CT) value of this PCR sample was about 20. |
The PCR protocol was run with different concentrations of template copies varied from 10 up to one million. Calculated parameters of x0 were plotted versus the number of templates showing the PCR standard curve.
As mentioned above, after thermal cycling of the PCR device a melting curve analysis28–30 was performed in order to determine the purity of the PCR. The fluorescence signal was approximated by a modified sigmoidal function:
![]() | (4) |
The fitting error shows only a small difference between the measured data and the fitting curve, showing that there is only one PCR product with a limited amount of by-products. The purity of products was proven by the results of capillary electrophoresis (see inset in Fig. 5).
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Fig. 5 Fluorescence signal from melting curve analysis (black line hidden behind the red line) approximated by a sigmoidal function shown in eqn (4) (red line). The deviation between measured data and the sigmoidal function is negligible demonstrating the purity of the PCR product. The negative value of its derivation is shown in blue. Results of capillary electrophoresis (see the inset) demonstrate the PCR product purity. |
The most important contribution of this work is the development of a system, which enables the micromachined chip to be separated from the disposable glass slide where the PCR takes place. The disposable part of the PCR system completely eliminates the danger of sample to sample cross contamination. As the glass does not require any processing except for the cleaning and fluorosilane coating, it costs just a few pennies. The thermal properties of the system are dominated by the sample volume, which can be further reduced making the PCR even more economical.
The whole system, as described above, can be mass-produced using well-established micromachining processes. We are currently developing a battery operated system: the PC will be replaced by a single chip processor, the fluorescence microscope and PMT detector will be substituted by a blue light emitting diode (LED), a miniaturized optical unit containing a photodiode and the optical signal processing will be integrated at the PCB with the microchip controller. Our goal is to complete the development of a truly economical and portable lab-on-a-chip system.
We would also like to thank the two anonymous reviewers who by their comments helped us to improve this paper.
Footnote |
† Elements SHELL-57 with real constants of 450 µm for silicon and 170 µm for glass were used. |
This journal is © The Royal Society of Chemistry 2006 |