performing microchannel temperature cycling reactions using reciprocating reagent shuttling along a...

Performing microchannel temperature cycling reactions using reciprocatingreagent shuttling along a radial temperature gradient

Ji-Yen Cheng,* Chien-Ju Hsieh, Yung-Chuan Chuang and Jing-Ru Hsieh

Received 21st January 2005, Accepted 4th April 2005

First published as an Advance Article on the web 22nd April 2005

DOI: 10.1039/b501061f

This study develops a novel temperature cycling strategy for executing temperature cycling

reactions in laser-etched poly(methylmethacrylate) (PMMA) microfluidic chips. The developed

microfluidic chip is circular in shape and is clamped in contact with a circular ITO heater chip of

an equivalent diameter. Both chips are fabricated using an economic and versatile laser scribing

process. Using this arrangement, a self-sustained radial temperature gradient is generated within

the microfluidic chip without the need to thermally isolate the different temperature zones. This

study demonstrates the temperature cycling capabilities of the reported microfluidic device by a

polymerase chain reaction (PCR) process using ribulose 1,5-bisphosphate carboxylase large

subunit (rbcL) gene as a template. The temperature ramping rate of the sample inside the

microchannel is determined from the spectral change of a thermochromic liquid crystal (TLC)

solution pumped into the channel. The present results confirm that a rapid thermal cycling effect

is achieved despite the low thermal conductivity of the PMMA substrate. Using IR thermometry,

it is found that the radial temperature gradient of the chip is approximately 2 uC mm21. The

simple system presented in this study has considerable potential for miniaturizing complex

integrated reactions requiring different cycling parameters.

1 Introduction

The polymerase chain reaction (PCR) technique represents a

major advance in modern life sciences. This technique has been

applied to a diverse range of research fields, including

molecular biology, forensic science, and genomics. In recent

years, much effort has been spent in miniaturizing the PCR

technique in order to enhance reaction throughputs while

simultaneously minimizing reagent consumption.1–5 It is anti-

cipated that the successful miniaturization of the PCR process

will enable the extension of this particular technique to a

broader range of clinical analysis processes.

The published literature contains various references relating

to the construction and configuration of miniaturized PCR

devices. For example, studies have addressed different chip

substrate materials and PCR devices with various reaction

vessel format.1,2,6 Briefly, reaction vessels are categorized as

having either a microchannel format or a micro-well format.

Regarding the substrate materials, glass, quartz and silicon (Si)

represent the most commonly employed inorganic substrates,

while organic substrates and reaction vessels are typically

fabricated using poly(tetrafluoroethylene) (PTFE), poly-

carbonate (PC), or poly(dimethylsiloxane) (PDMS). Some of

these substrates provide good optical transparency, while

others have good thermal conductivity.

Many reports have been presented describing different

temperature cycling approaches. There are two basic tempera-

ture cycling strategies. The first strategy involves the direct

heating and cooling of the entire reaction vessel by means of a

temperature-controlled heater and cooler. Typically, strategies

of this type adopt either contact4,7,8 or non-contact9–12

heating/cooling approaches. In the second strategy, sometimes

referred to as spatial-temperature cycling, the sample is heated

and cooled by flowing the sample through a designated

sequence of temperature-controlled blocks.4,13,14

The discrete micro-well format4,8–10,13,14 has the advantage

that the number of temperature cycles can be varied while the

same chip design is used. In addition, extending a single-well to

a multiple-well reaction chip in order to increase throughout is

straightforward15 and can be achieved simply by arranging

multiple reaction wells in an array format. By contrast,

adjusting the number of temperature cycles is more proble-

matic when using a flow-through microchannel reaction chip,

i.e. the microchannel format, since the number of cycles is

determined by the physical configuration of the device.

Although the closed-loop microchannel PCR device reported

in ref. 13 has circumvented this disadvantage, it is nevertheless

necessary to increase the number of heating blocks when

additional temperature zones are required. Moreover, in the

flow-through microchannel format, each sample occupies a

greater chip area than in the well format. Therefore, multi-

plexing the microchannels of flow-through PCR chips greatly

increases the necessary chip size. However, since the heat

exchange zones of the flow-through microchannel PCR format

are maintained at preset constant temperatures, it is not

necessary to heat and cool the reaction chamber repeatedly

during the temperature cycling process. Therefore, tempera-

ture cycling in a flow-though microchannel device is more

efficient than in devices with a micro-well format. However,

incorporating additional temperature zones complicates the

design of flow-through chip. For example, a recent report

integrating reverse-transcription with PCR has increased the*[email protected]

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This journal is � The Royal Society of Chemistry 2005 Analyst, 2005, 130, 931–940 | 931

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number of heating block from three to four in the micro-

channel chip.14

The objective of the current study is to develop a tem-

perature cycling reaction microchip that combines the benefits

of the several microchip PCR formats described above. That

is, the efficient temperature cycling can be performed, as is in

the flow-through microchannel PCR chip, while the flexibility

of varying the cycle number and varying the number of

temperature zones is maintained, as is in the micro-well PCR

chip. Very recently, a chip PCR scheme that shuttles sample

back and forth in a microchannel has been reported.16,17

An annular or circular continuous temperature distribution

was generated in the reported microfluidic chip and the liquid

sample plug was pumped back and forth along the radial

direction of the chip to achieve rapid temperature cycling.

Since a continuous temperature distribution is sustained in the

chip, the number of temperature zones is not limited. Although

a similar spatial temperature cycling strategy has been utilized

previously in the literature,4,13,14 the temperature zones have

been distributed discretely rather than continuously as in the

present case.

This study integrated a microfluidic chip fabricated from

poly(methylmethacrylate) (PMMA) with a resistive heater chip

fabricated from transparent indium tin oxide (ITO)-coated

glass. The two chips were both machined by CO2 laser, which

enables the chip to be rapidly re-designed and re-fabricated as

required. The integrated chip/heater assembly provides an

economic and flexible means of generating a radial tempera-

ture gradient for temperature cycling reactions.

In this study, non-contact IR thermometry was utilized to

assess the radial temperature gradient on the chip surface. The

distribution symmetry of the temperature distribution was

then quantitatively examined. Meanwhile, the literature has

reported several methods for internal temperature measure-

ment, including fluorescence quenching,18 fluorescent mole-

cular beacons,19 and thermochromic liquid crystal (TLC).20 In

this study, the internal temperature of the microchannel was

calibrated on the basis of the spectral change of a TLC

solution pumped through the microchannel to monitor the

temperature ramping rate inside the microchannel.

Although many suitable candidates exist for the microfluidic

chip substrate material, there are several advantages of

adopting PMMA. For example, PMMA is transparent, and

hence enables the application of conventional optical detection

methods when evaluating the microfluidic chip’s heating

characteristics.21 Furthermore, PMMA is an economic mate-

rial. Using the CO2 laser micromachining technique,22,23

economic and versatile microfluidic PMMA-based chips

can be fabricated in a matter of minutes. This particular

fabrication approach offers great flexibility for microfluidic

chip development. PMMA also has favorable biocompatibility

properties and has been used for many different biomedical

applications24 in recent decades. It has been reported that

protein adsorption on PMMA is less than on other forms of

plastics.25–27 Particularly, the adsorption on PMMA is less

than that which occurs on polycarbonate (PC) and polystyrene

(PS) for contact times of less than 30 min.25 Some widely used

substrate materials such as poly(dimethylsiloxane) (PDMS)

and silicon (Si) are known to result in substantial protein

adsorption,3,28 which may cause the PCR amplification

process to fail.29 Moreover, microchannels in Si chips can

not be integrated with photo-detection methods involving the

use of optical paths perpendicular to the substrate plane

because Si is not transparent to visible light. Glass and quartz

are also frequently used as microfluidic chip substrate

materials.30–32 However, the fabrication processes required

for these substrates are not as convenient as those for plastics.

Accordingly, PMMA was chosen as the substrate material of

the developed microfluidic device.

Addressing a perceived lack in the literature, this study

documents the performance of PMMA as a microchip PCR

substrate material. This study also examines the temperature

equilibrium rate attained in the reported device and confirms

that the design facilitates a rapid heat transfer from the chip to

the liquid sample. Although the present study takes the PCR

process as an example to demonstrate the temperature cycling

process, the execution of other applications requiring addi-

tional temperature zones can also be accomplished very easily

using the reported device.

2 Experimental procedure

2.1 Design and fabrication of heater chip and microfluidic chip

The reported reaction system comprises two basic compo-

nents, namely a heater chip and a microfluidic chip. The heater

chip is fabricated from ITO-coated glass, while the micro-

fluidic chip is fabricated from a PMMA sheet. In the reported

design, the two components are clamped firmly together to

ensure a close contact between them.

Since ITO is a transparent conductor, its use enables the

integration of an optical detection technique through the chip

plane. A deposited ITO film has been used as the chip heater

material in a previous study.32 However, the present study

differs in that it uses a piece of commercially available ITO-

coated glass (243735X0, Merck) as the heater substrate. In the

current fabrication process, a circular chip of diameter 85 mm

was cut from the ITO glass blank. It is noted that this diameter

was chosen merely to enable convenient manual handling. In

practice, the diameter can be modified to suit a particular

application. The thickness of the ITO film was y80 nm.

Discrete ITO strips were prepared on the chip surface to serve

as resistive heaters by scribing the ITO glass surface using a

CO2 laser scriber (M-300, Universal Laser). The sheet

resistance of the ITO film was 22.5 ohm per square (V %21).

Since the ITO film has a uniform thickness, the resistance of

each strip is governed by its width and length. In general, the

position, geometry, and resistance of each resistor strip can be

designed using computer software such as CorelDraw or

AutoCAD. As an example of the typical power consumption

of the heating device, a 3 mm wide circular strip with radius of

30 mm (hence the resistance of 2025 V) has heating power of

3.5 W. The maximal temperature generated by this heating

ring is y100 uC, as shown in the result below. The positional

accuracy of the scriber used in the present study was 10 mm

and the minimum attainable width of the ITO strip when using

the CO2 laser scriber was approximately 100 mm. In this study,

the ITO heating strips (indicated by the shaded regions in the

ITO chip in Fig. 1) were ring-shaped in order to establish a

932 | Analyst, 2005, 130, 931–940 This journal is � The Royal Society of Chemistry 2005

radial temperature gradient. As shown, two concentric heating

rings, connected in parallel by a pair of wide ITO strips of low

resistance, were designed on the chip. During fabrication, the

location of each resistor strip was specified by the computer

drawing software used during the scribing process. Clearly,

when a constant voltage is applied to the ITO strips via a pair

of conductive pins, a lower resistance results in a higher power,

and vice versa. When the heating ring with a lower power

was positioned near the center of the chip and the ring with

a higher power was located near the circumference, as with

the ITO pattern shown in Fig. 1, a circular bowl-shaped

temperature distribution was generated (see below, Fig. 3(b)).

Alternatively, positioning the high-power heating ring in the

center and the low-power ring near the circumference

produced a cone-shaped temperature distribution (Fig. 3(a)).

In summary, in the reported microfluidic device, the ITO

resistors can be dimensioned and positioned at an appropriate

pre-determined location to obtain the required temperature

distributions. Alternatively, a similar effect can be obtained

without modifying the chip design but simply by modifying the

voltages applied to the two ring heaters.

The use of a CO2 laser for PMMA surface machining has

been reported previously.22,23 The CO2 laser micromachining

of polycarbonate (PC) has also been studied.33,34 Using these

techniques, trench widths as low as y80 mm can be rapidly

prepared on plastic surfaces. In the present study, the required

trench pattern was designed using commercial computer

software and a CO2 laser scriber was then used to etch the

required trenches onto the PMMA substrate (thickness 5 2 mm,

Yen-Nan Acrylic Co. Taiwan). On-chip via holes for sample

loading were also etched into the chip using the laser scriber.

Various trench widths and depths were obtained by controlling

the laser beam diameter and the laser power. A higher power

was used to cut a circular PMMA chip of the same diameter as

the heater chip (85 mm) from a PMMA sheet. The trenched

PMMA chip was then thermally bonded with a blank PMMA

substrate (a cover plate) to form a chip with sealed channels, as

described in our previous study.22

The PMMA chip drawn in Fig. 1 shows the microfluidic

channel configuration used to perform the PCR process. It can

be seen that the overall configuration comprises a total of three

interconnected arcs. In this particular example, the inner arc is

used to perform the annealing step, the intermediate arc is for

the extension step, and the outer arc is for the denaturing step.

Although the minimal feature attainable by the current laser

scriber was as low as y80 mm, the microchannel in this study

was intentionally enlarged to increase the volume capacity so

that enough sample was collected for gel electrophoresis

analysis. The channel width and depth were specified as 1.2

and 0.6 mm, respectively. It was found that the laser ablation

process yielded a smooth and transparent channel wall.

Previous studies in the literature have employed a CNC milling

to fabricate channels of similar dimensions.25,35 However, the

current CO2 laser scribing process is suitable for the machining

of both the ITO glass used for the heater chip and the PMMA

used for the microfluidic chip substrate. The use of a single

machining process simplifies the overall fabrication procedure

considerably. In the fabricated PMMA microfluidic chip, the

channel length is approximately 70 mm. However, the channel

length can be easily re-fabricated using the CO2 laser scriber.

After bonding the trenched PMMA chip with the cover

plate, an adaptor was glued to the sample-loading hole on the

chip to facilitate the attachment of a microtube (OD 1.6 mm).

The microchannel walls were then coated with amorphous

Teflon (PFC802A, Cytonix) to reduce their surface energy in

order to ensure that the liquid sample would form a unified

plug when pumped through the channel. The microtube was

then connected to a computer-controlled syringe pump. It is

noted that the microchannel had only one opening for sample

loading. Hence, connecting this opening to the syringe pump

sealed the entire channel.

To heat the microfluidic chip using the heater chip, the two

chips were stacked together and held securely in place by

means of clamps and screws to ensure a good thermal contact.

A power supply is used to provide a stable voltage to the

heater chip.

Fig. 1 Microfluidic system setup showing assembly of microfluidic chip and heater chip. The two chips are clamped together and secured by

screws. All components other than the glass chip and the conductive pins are made of plastic. The heater chip is fabricated from ITO-coated glass.

The shaded area denotes the energized strips used for heating. The microfluidic chip is fabricated from PMMA. Note that only one opening exists.

In this example, three interconnected arcs are fabricated at three different radii. The radius for both chips is 85 mm.


Following sample loading, a syringe pump was connected to

the opening and air was pumped into the channel. In the current

design, the injection of compressed air pushes the liquid plug

forward into the microchannel. Conversely, when the pressure

of the compressed air in the syringe pump is reduced, the liquid

plug is pulled back towards the sample loading opening.

Hence, the liquid plug can be positioned in the desired tempera-

ture zone of the microchannel simply by controlling the

position of the syringe piston. The plug can then be shuttled

back and forth between the desired temperature locations by

varying the syringe pressure appropriately. The actual pressure

increase is calculated by the plug position. Note that since in

the current microfluidic device, the temperature gradient is

radial and continuous, the local temperature of any region can

be deduced from its radial position on the chip.

2.2 Measurement of chip surface and microchannel temperatures

To eliminate perturbations caused by the use of contact-type

thermal probes, this study employed a non-contact IR

thermometry technique to assess the surface temperature of

the chip. A mid-IR (8–12 mm) two-dimensional temperature

distribution thermal image was acquired by an IR camera

(120 6 120 pixels, IR SnapShot, Model 525, IRCON). The

temperature measurement accuracy of the IR camera was 2 uCand the spatial resolution of the IR image was 0.9 mm.

However, the internal temperature of the chip can not be

obtained using the same IR thermometry approach. Conse-

quently, in the present study, this temperature was determined

by the spectral change of a TLC solution passed through the

channel. Two different TLC fluids were used, one with a color

changing temperature of 71 uC (blue to transparent, band-

width 3 uC, Product No. PD Blue 71, New Prismatic Enterprise,

Taiwan) and another with a color changing temperature of

60 uC (black to transparent, bandwidth 3 uC, Product No. PD

Black 60, New Prismatic Enterprise, Taiwan). In both cases,

the TLC was suspended in water with 1% Tween 20 (X251-07,

J. T. Baker) added to enhance dispersion. The concentration of

the TLC solution was 5% w/v.

The static internal temperature was calibrated using the

following procedure. Initially, the TLC solution was flowed

into a microchannel having two openings (Fig. 2). Once

thermal equilibrium conditions had been achieved (y10 s), the

TLC color in the regions where the temperature exceeded the

color changing temperature became transparent. Meanwhile,

the color of the TLC in the regions of the microchannel having

a temperature lower than the color changing temperature

remained unchanged, as shown in Fig. 2. Finally, the position

of the microchannel at which color transition took place

corresponded to the point at which the local temperature was

the same as the TLC operation temperature.

2.3 Temperature ramping rate in microchannel

A TLC solution was also used to determine the temporal

temperature change in the microchannel. In determining the

temperature ramping rate, the measurement setup was the

same as that shown in Fig. 2 and a fiber spectrometer

(USB2000, Ocean Optics) was used. As described above, the

chip incorporated a microchannel with two openings, one of

which was used to introduce the reagent and another which led

to the waste buffer. Initially, a syringe was used to introduce

the TLC solution until the entire microchannel had been filled.

In the microchannel regions having a temperature lower than

71.5 uC, the liquid crystal appeared blue in color (dark liquid

in Fig. 2). The low-temperature absorption was measured by

the fiber spectrometer and the maximum absorption wave-

length was then determined for temporal detection. The

maximum absorption wavelength was found to be 600 nm.

Once temperature equilibrium conditions had been achieved,

the visible absorption of the liquid crystal inside the micro-

channel was continuously monitored by the fiber spectrometer

aligned with the region of the microchannel within which the

TLC was transparent and was located at a distance of approxi-

mately 1 mm from the color transition point, as shown in

Fig. 2. At the beginning of the temporal temperature change

measurement process, 200 ml of liquid crystal at room tempera-

ture was pumped through the channel at a flow velocity of

y2 m s21. The liquid flow was then stopped and the liquid

maintained in a fixed position aligned with the fiber head of the

spectrometer. The liquid crystal absorption increased signifi-

cantly when the TLC solution was initially pumped through the

channel. However, the absorption then decayed to its initial

value when the liquid crystal became transparent. The temporal

change of the absorption in the range 600 ¡ 30 nm was

integrated, recorded and digitized by the spectrometer. The

temporal resolution was set to be 0.3 s. It is noted that the volume

of the TLC solution consumed in the temporal measurement

(200 ml) process was much larger than that used for the

temperature cycling reaction process (4.5 ml), which ensured a

complete heat transfer for sample volumes smaller than 200 ml.

2.4 Conventional PCR, microchannel PCR and gel

electrophoresis

The PCR cocktail consisted of 100 pg ml21 ribulose 1,5-

bisphosphate carboxylase large subunit (rbcL) template DNA,

Fig. 2 Setup for measuring TLC spectral change in microchannel. A

miniature fiber spectrometer is used to monitor the static absorption

and the temporal absorption change of the TLC solution. Cold TLC

(dark liquid in the syringe) is pumped into the microchannel at a speed

of y2 m s21.


0.16 mM of each dNTP base (1969064, Roche Boehringer

Mannheim Indianapolis, IN, USA), 10 mM betaine (20424-

1000, Acros, Morris Plains, NJ, USA), 0.2 mM of each primer

(forward primer: 59-GCGTGTTGTTGGAGAAAAAGA-39;

reverse primer: 59-GGACGACCATACTTGTTCAATTT-39),

5 mg ml21 bovine serum albumin (BSA; B-4287, Sigma, St.

Louis, MO, USA), 0.2 units ml21 FastStart taq (2032953,

Roche Boehringer Mannheim), and a reaction buffer (50 mM

Tris-HCl, 10 mM KCl, 5 mM (NH4)2SO4 and 2 mM MgCl2,

pH 8.3 at room temperature). The anticipated product length

was 240 bp, calculated by the primer alignment to the template

sequence (Accession number: Nc001879). The reaction cocktail

was loaded into the microchannel manually using a micro-

pipette. The temperature cycling process was then performed

using the procedure described below. Conventional PCR is

performed on a commercial PCR cycler (Bio-Rad, MyClear

Thermal Cycler). The total volume is 100 ml.

The concept of temperature cycling using the preset

temperature gradient involves pumping the sample plug along

the microchannel and shuttling it back and forth between

microchannel regions having the desired temperatures. In the

PCR process, 4.5 ml of the reaction mixture was loaded into the

well and then pumped into the channel. Surface tension effects

acting on the aqueous sample caused the formation of a short

unified liquid plug in the microchannel. It was found that

without the Teflon coating in the inner wall of the

microchannel, the microchannel PCR failed. The specific

reason was not investigated in this study. A syringe pump

was used to pump air into the channel, thereby driving the

PCR sample plug into the microchannel. In the initial

denaturing step, the PCR sample plug was retained in the

denaturing arc (95 uC) for 180 s. The subsequent temperature

cycling steps were as follows. The sample plug was held in the

annealing arc (57 uC) for 30 s and then pumped to the

extension arc (72 uC) and incubated for 30 s. The plug was then

pumped to the denaturing arc and incubated for 30 s. To

complete the cycle, the sample plug was moved back to the

annealing arc. The pumping speed is about 10.8 mm s21 (flow

rate 5 9.7 ml s21). Note that both the pumping speed and the

number of temperature cycles were controlled by the program

driving the syringe pump. Following temperature cycling, 2 ml

of the product was retrieved from the via hole by the micro-

pipette and incubated with 1 ml of loading dye, Bromophenol

Blue (B-5525, Sigma, St. Louis, MO, USA), and fluorescence

dye, SYBR Green I (S-7585, Molecular Probes, Engene, OR,

USA), for 15 min under dark conditions. A DNA ladder

(Invitrogen) was treated with the same dyes for comparison.

The PCR products were then analyzed by 2% agarose gel

electrophoresis in 0.5 6 TAE (Tris/acetate/EDTA) buffer.

3 Results and discussion

3.1 Radial surface temperature distribution of microfluidic chip

In investigating the radial surface temperature distribution of

the microfluidic chip, the chip was securely clamped to the

heater chip and then heated. The radial temperature distribu-

tion was evaluated on the basis of the observed chip surface

temperature. An IR camera was used to measure the chip

surface temperature. The IR thermometry method deduces the

substrate temperature by measuring the IR emissions (mid-IR,

wavelength 5 8–12 mm) from the substrate. Hence, a limitation

of this technique is that only surface temperatures can be

obtained since most materials are not transparent to mid-IR

emissions.

Depending on the relative geometrical positioning of the

high-power and low-power heating rings, two different types

of radial temperature distribution can be established on the

chip. When the high-power heating ring is placed near the chip

center, a cone-shaped temperature distribution, as shown in

the three-dimensional plot in Fig. 3(a). Alternatively, an

annular type temperature distribution is obtained when the

high-power heating ring is positioned near the circumference

of the chip, as shown in Fig. 3(b). It is noted that the present

design does not involve the use of active cooling devices to

maintain the temperature gradient. Rather, the distribution is

sustained by natural thermal conduction mechanisms. The

omission of a cooling device enables a reduction in the

footprint of the completed chip assembly.

The two spatial temperature distributions shown in Fig. 3(a)

and (b) are both circularly symmetric. However, it can be

seen that there exists a non-symmetric sector subtending

an angle of approximately 40 degrees at the center of the

chip. This deviation from the radial symmetry is caused by the

non-symmetric hardware structure in this region of the chip

comprising of the probe pins and the two ITO strips

connecting the two heating rings. The non-symmetric region

can be minimized by adopting an optimized heating coil

pattern. Alternatively, a non-contact inductive heating

strategy12 can be applied to eliminate the use of the contacting

pins and connecting ITO strips. In the current study, this

problem is circumvented by specifically designing the micro-

fluidic channel within a region of the chip positioned away

from this non-symmetric sector.

The distribution symmetry can be quantitatively expressed

in terms of the roundness of the isothermal zone. Using the

data presented in Fig. 3(b), IR image pixels corresponding to a

surface temperature ranging from 55 to 57 uC are analyzed.

The corresponding region is an annular temperature zone. The

roundness of this zone can then be examined. The radius

values around the isothermal circumference are found to be 23

pixels with a standard deviation of 0.37 pixels. Since the IR

imaging resolution is 0.9 mm/pixel, the corresponding spatial

deviation is 0.33 mm. These results indicate that the roundness

of the isothermal region is highly satisfactory. The deviation is

smaller than the image resolution of the IR imager. At this

radius, the temperature gradient is 2 uC/pixel (based on the

data in Fig. 3(b)) and therefore, the temperature uncertainty is

0.74 uC (2 uC/pixel 6 0.37 pixel). This analysis takes account

of the non-symmetric sector described above and indicates

that the temperature distribution asymmetry effect appears

mainly in the high-temperature regions. It is therefore reason-

able to estimate the surface temperature of extension and

annealing regions of the chip on the basis of its radial distance

from the center.

Several significant characteristics were observed in the

developed microfluidic system. For example, the radial

temperature distribution discussed above indicates that the

reported simple design is capable of generating a large


temperature span. The temperature ranges of the temperature

distributions shown in Fig. 3(a) and (b) are found to be 25 uCto 81 uC and 47 uC to 82 uC, respectively. It is noted that these

values refer to surface temperatures rather than internal

temperatures. Expanding the range to a higher temperature

domain can be achieved by using a suitable heat resistant

substrate material, such as PC or glass. Furthermore,

additional cooling devices such as annular thermoelectric

coolers can be integrated with the reported device to extend its

operating range towards the lower temperature domain. The

miniaturized microfluidic heater/cooler presented in Guijt’s

work19 can also be used without any significant increase in the

device size. Additionally, in the reported design, the tempera-

ture changes gradually and continuously along the radial

direction. Hence, there is no need to adopt thermal isolation

structures such as those micro-structures reported in some

delicate temperature cycling chips8 to isolate the different

temperature zones. Furthermore, the temperature at any

local chip position can be predicted solely on the basis of its

radial position. Therefore the reported design allows for

the incorporation of additional temperature zones without

the need to add further heating elements. For example, the

telomerase activity assay (i.e. Telomerase-PCR-ELISA, or the

TRAP assay),36 which requires two additional temperatures of

30 uC and 37 uC, can be integrated with PCR simply by driving

the sample to the desired temperature zones. Additionally, the

sample temperature overshooting phenomenon reported for

some micro-well PCR chip33 does not occur when the sample is

pumped correctly through the desired temperature zones.

Moreover, multiple sectors can be arranged on the same chip

in order to perform either independent or integrated reactions.

These sectors can be assigned identical or independent

temperature cycling parameters as required. For example, the

effects of different temperature cycling parameters, including

the incubation temperature, the incubation time, and the

temperature ramping rate, can all be tested using a single

microfluidic chip.

In order to utilize the continuous temperature gradient for

the execution of microfluidic temperature cycling reactions, it

is necessary to obtain the temperature distribution inside the

microchannel in addition to the surface temperature distribu-

tion discussed above. It is anticipated that the internal tem-

perature distribution will also maintain a radial symmetry,

although the gradient may vary slightly from that of the

surface temperature distribution.

3.2 Microchannel interior temperature distribution

When detecting the temperature of an object using a detection

probe, the probe absorbs heat from the object and conducts

heat away from it. Hence, the original temperature is

perturbed and hence can not be measured accurately. The

use of thermochromic liquid crystals (TLCs) provides an

appropriate alternative for internal temperature measure-

ment.33,37 This technique is highly effective in identifying the

Fig. 3 3-D plots of surface temperature distribution of circular

PMMA microchip. (a) Shows a cone-shaped temperature distribution.

The temperature varies from 81 to 25 uC from the center to the

circumference. (b) Shows bowl-shaped temperature distribution. The

temperature varies from 82 to 47 uC from the rim to the center. In (a)

and (b), a brighter color indicates a higher temperature. (c) Shows the

internal microchannel temperature profile, calculated by eqn. (2), of

cone-shaped and bowl-shaped temperature distribution along the chip

radius. The range for the cone-shape type extends from 25 to 101 uC,

while that for the bowl-shape type extends from 55 to 102 uC.


regions having the same temperature as the TLC’s operation

temperature. In measuring the internal temperature distribu-

tion, this study first determines the temperature distribution

by IR thermometry and then correlates the internal tempera-

ture to the surface temperature by means of a thermo-

conducting model calibrated by the TLC internal temperature

measurement.

In developing the thermo-conducting model, the thermal

conductance through the chip is first simplified into the form

of a one-dimensional thermo-resistor system.38 This system

can then be transferred to the equivalent thermal circuit. In

this circuit, the heat transfer rate is expressed as:

qx~Ts{T?

1=hA~

Tc{Ts

L=kA(1)

Hence:

Tc~Lh

kz1

� �Ts{T?ð ÞzT? (2)

where Ts is the surface temperature of the PMMA microfluidic

chip, Tc is the temperature in the channel inside the chip,

and T‘ is the ambient temperature. Furthermore, L is the

thickness of the trenched PMMA chip, k denotes the thermal

conductivity of PMMA, and h is the convection heat transfer

coefficient of air. Finally, A denotes the area of the PMMA

surface perpendicular to the direction of heat transfer. Since

the model simplifies the heat conduction to a one-dimensional

form, the approximation is valid at positions distant from the

chip circumference.

Eqn. (2) allows the temperature inside the channel of the

PCR chip to be estimated on the basis of the measured surface

temperature. However, it is important to compare the

calculated channel temperature to that indicated by the TLC

method. Initially, a TLC solution with an operation tempera-

ture of 71.5 uC is used to identify the region of the

microchannel having the equivalent internal temperature.

The distance from this point to the circumference is found to

be 21 mm. This position is reasonably distant from the chip

circumference. The surface temperature at this point is then

established from the IR thermometry image and the corres-

ponding internal temperature is calculated from eqn. (2) as

Tc 5 73.5 uC. This temperature corresponds very well to that

indicated by the TLC approach (71.5 uC); particularly when it

is considered that the TLC color is known to deviate from its

original hue under the influence of ambient light.37 Therefore,

this study uses eqn. (2) in conjunction with the IR thermo-

metry technique to deduce the two-dimensional internal

temperature distribution.

3.3 Reciprocating temperature cycling

The following paragraph describes the use of the two-

dimensional temperature distribution in the temperature

cycling process. The second type of heater design is taken

as an example. In this design, the temperature distribution

exhibits a high-temperature rim near the chip circumference

(Fig. 3(b)). The surface temperature of this high-temperature

zone is 75 uC and the calculated channel temperature is 102 uC.

This temperature is slightly higher than the glass transition

point of the PMMA substrate. However, no substrate

deformation is observed in the current short-term applications

(,40 min). Inside the high-temperature rim, the temperature

decreases with decreasing radius. In this example, the lowest

channel temperature is 55 uC. As described above, the internal

channel temperature can be calculated from the measured

surface temperature by eqn. (2). The internal temperature

profile along the chip radius is shown in Fig. 3(c) for both

heater configurations. Utilizing the self-sustained temperature

gradient, it can be seen that the reported chip is capable of

providing temperature zones ranging continuously from 55 to

y100 uC.

Using the radially symmetric temperature gradient, tem-

perature cycling can be achieved by moving the sample

backward and forward along the chip radius. Since the local

temperature can be calculated on the basis of the distance from

the chip center, the microfluidic channel can be designed such

that it passes through the required temperature zones. When

the liquid sample is pumped into the channels and moved to

the desired temperature zones, it assumes the local temperature

via heat transfer.

Since the microfluidic chip and the heater chip are fabricated

separately, various channel designs can be fitted to the same

heater chip to obtain different temperature cycling parameters.

Alternatively, different temperature cycling parameters can be

obtained without modifying either chip by means of an

appropriate sample pumping scheme. Under the latter

approach, when performing the temperature cycling reaction

process, the sample plug can be reciprocated at different

pumping speeds in order to change the temperature ramping

rate. The sample plug can also be maintained in a stationary

position so that it is incubated at the constant temperature

associated with that particular location.

3.4 Temperature ramping rate in microchannel

It is important to examine the maximum temperature ramping

rate of the sample in order to enhance the control of the

temperature cycling parameters. As discussed previously, the

temperature change inside the microchannel can not be

measured by IR thermometry and is therefore monitored by

the temporal spectral change of the TLC.

TLCs are a class of organic compounds that change their

light reflectivity characteristics with temperature. Although the

temporal color change can be recorded by a video camera, it is

more conveniently monitored spectrally since this avoids hue

interference caused by the light source. The temporal absorp-

tion change is easily recorded and digitized by a miniature

spectrometer.

Fig. 4 shows the abrupt absorption increase which occurs

when 200 ml of the TLC solution is pumped through the

detection point at a rapid velocity of 2 m s21. The absorption

then decays gradually as the temperature of the TLC rises

and its color becomes transparent. As shown, the temperature

of the liquid inside the microchannel reaches equilibrium con-

ditions at 71.5 uC in approximately 10 s from room tem-

perature (25 uC). The corresponding temperature ramping rate

is y4.5 uC s21. PMMA, as with other plastics, is recognized

as a poor thermal conductor. Its thermal conductivity is


0.193 W m21 K21.39 However, in the present study, temperature

equilibrium is attained very rapidly, which indicates that a

rapid heat transfer to the sample takes place inside the micro-

channel. The surface to volume ratio of the channel with the

width of 1.2 mm and the depth of 0.6 mm is 8 mm2 ml21. Hence,

the rapid heat transfer to the sample is reasonably attributed to

the large surface to volume ratio of the microchannel.

Although in the experiment above, temperature equilibrium

is attained after approximately 10 s, the time required for the

actual temperature cycling process is far less when a smaller

sample volume is used. An almost instantaneous color

bleaching was observed when the TLC was pumped through

the heated channel at a slow speed of 10 mm s21 (data not

shown). There are several reasons for this. First, the

temperature equilibrium observation in Fig. 4 is performed

by pumping a large volume of TLC solution at room

temperature into the channel in a very short time (i.e. 200 ml

in y0.3 s). This procedure is adopted in order to ensure that

the temperature of the TLC solution at the spectrum detection

point rises from room temperature so that color change can be

recorded. The heat required to raise the TLC temperature is

provided solely by the channel wall at the detection point and

the limited heating area prolongs the temperature rise time.

However, when sample is pumped through the microchannel

at low speed and the temperature cycling is performed slowly,

the temperature equilibrium time is far shorter. The estimation

of the fastest temperature ramping rate is as follows. Assuming

the sample is pumped at 10 mm s21 from the chip center to the

circumference, for which the corresponding temperature

change is from 50 to 100 uC, and the chip radius is 43 mm,

the corresponding temperature ramp rate is y12 uC s21. It can

be concluded that the large surface to volume ratio and the

absence of a large heat mass yield a rapid temperature cycling

of the sample in the microchannel.

3.5 Sample evaporation inside channel

Sample evaporation is often problematic in microfluidic

applications because the sample volumes are very small.

Hence, this study devised a sample-pumping scheme to reduce

sample loss during the temperature cycling process. The

reported method significantly reduces the sample evaporation,

particularly for samples processed in high-temperature zones.

Under standard atmospheric pressure, water evaporation is

vigorous at 100 uC. In order to avoid sample loss through

evaporation in the high-temperature zones during the tem-

perature cycling process, this study intentionally increases the

internal pressure when the sample plug is pumped to high-

temperature zones. In the reported approach, a single opening

is fabricated for each microfluidic channel. This opening serves

for the loading of the sample and for connection to the syringe

pump. When the opening is connected to the pump, the

microchannel is completely sealed by the syringe. Air is then

compressed into the channel in order to drive the sample plug.

The air behind the distal end of the plug is also compressed

when the plug is pushed forward. Since the opening is

deliberately positioned in the low-temperature region, the

channel pressure increases as the sample plug is driven toward

the distal end of the channel in the direction of the high-

temperature zone. The internal pressure can be calculated by

Boyle’s law. Alternatively, it can be determined simply by

dividing the cross-sectional area of the syringe pump into the

force required to push the sample plug (Pressure 5 Force/

Area). In the present example, the pressure is calculated to be

y7 kg cm22 when the sample is driven to the high-temperature

zone. The water vapor pressure is 1/7 that of the compressed

air. At this pressure, water boils at a temperature of 166 uC.40

Therefore, the sample evaporation is greatly reduced where the

local temperature is only around 100 uC. As shown in Fig. 5, in

the current PCR temperature cycling example, the reduction in

sample volume varies as a function of the number of PCR

cycles. It can be seen that y70% of the sample remains as a

unified plug following 30 temperature cycles. The addition of

betaine is at low concentration (10 mM) compared to that in

the literature (1–2.5 M,41) and does have significant effect on

boiling temperature. The evaporation reduction is therefore

considered to be resulted from pressure increase. Without the

high pressure achieved in the current design, the sample would

boil in the denaturing zone and would dry in seconds.

Fig. 5 Minimized sample evaporation inside microchannel vs.

number of temperature cycles. Evaporation is minimized by pressuriz-

ing the sample during heating.

Fig. 4 Temporal spectral change of TLC (color changing tempera-

ture 71.5 uC) solution inside microchannel. The color change is

measured by visible absorption spectroscopy. The inset shows the

absorption spectrum of the liquid crystal.


jycheng

反白

3.6 Microchannel PCR using radial temperature gradient

This study takes the well-known PCR process as an example

in order to demonstrate the successful temperature cycling

reaction achieved using the reported microfluidic device.

However, it should be noted that the same chip could also

be extended to the execution of processes other than PCR.

Fig. 6 presents the agarose gel electrophorogram of the

amplified PCR product. The first lane indicates the DNA

ladder, while the third lane is the product from the

microchannel PCR. It can be seen that the anticipated 240 bp

product amplified from the rbcL gene template has been

obtained. The reaction took a total of 38 min for the micro-

channel PCR with it is 3.5 h for the conventional PCR. The

injected sample volume was 4.5 ml and the amplification was

performed over 23 cycles. As with the well-based microchip

PCR,6,15,42–44 the cycle number can be adjusted as required in

the reported chip and is not limited by the original channel

configuration, as is the case for flow-through PCR chip.

Furthermore, the reported microfluidic chip can be divided

into a number of discrete sectors such that multiple channels

can be configured on the same chip. Approximately 10 PCR

channels can be fabricated on the current chip with a diameter

of 85 mm diameter. However, further optimization is currently

under development with the intention of increasing this

number in future versions of the chip. Currently with the

limited channel number but with the low power consumption

(3.5 W), the reported device may be useful as a portable device,

e.g. a lab-on-a-chip PCR machine.

4 Conclusion

This paper presented a simple and versatile PMMA micro-

fluidic system for temperature cycling reactions utilizing a self-

sustained radial temperature gradient. It has been shown that

this device can generate a large temperature span ranging

from room temperature to y100 uC without the need for a

cooling device or active temperature control mechanisms.

Furthermore, the temperature gradient can be maintained

without the need for complex temperature control electro-

nics.31 Furthermore, no thermal isolation structures are

necessary to separate zones of different temperature. The

completed chip assembly provides a powerful tool for

integrating biological and chemical reactions on a single chip.

The PMMA microfluidic chip and the ITO resistor heater

chip were both fabricated using an economic CO2 laser

scriber. The entire manufacturing process is straightforward

and highly versatile. The overall approach is eminently suitable

for low-cost rapid prototyping.

A non-invasive full-chip temperature characterization pro-

cedure has been presented. IR thermometry and TLC

spectral change observation techniques have been combined

to calibrate the temperature distributions on the chip surface

and within the microchannels, respectively. Based upon the

observed absorption change of the TLC, it has been shown

that the liquid sample in the channel attains a temperature

equilibrium condition in less than 10 s. For low sample volume

(y4.5 ml) the temperature ramping rate is estimated to be

y12 uC s21. It should be emphasized that a spectral method

rather than a colorimetric method is employed to monitor

the TLC color change. Hence, the potential interference of

ambient light on the TLC color is avoided.

A novel strategy for the temperature cycling process has

been presented, together with a scheme designed to reduce

sample evaporation. Although the microchannel PCR process

has been taken as an example to demonstrate the capability

of the chip to perform temperature cycling reactions, the chip

can readily be extended to other processing operations. The

reported scheme successfully combines the benefits of the well-

based PCR and continuous flow microchannel PCR devices on

a single chip.

Abbreviations

PCR, polymerase chain reaction; PMMA, poly(methylmetha-

crylate); ITO, indium tin oxide; PTFE, poly(tetrafluoroethyl-

ene); PC, polycarbonate; PDMS, poly(dimethylsiloxane).

Acknowledgements

This work is partially supported by the National Science

Council, Taiwan under the contract number NSC 93-2113-M-

001-037.

Ji-Yen Cheng,* Chien-Ju Hsieh, Yung-Chuan Chuang andJing-Ru HsiehResearch Center for Applied Sciences, 128 Sec. 2 Academia Rd., Taipei,11529 Taiwan E-mail: [email protected];Fax: 886-2-2782-6680; Tel: 886-2-2789-8000-34

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