RAPID DETECTION OF PATHOGENS AND THEIR ANTIBIOTIC SUSCEPTIBILITY
USING SIMPLE MICROFLUIDICS AND CCD IMAGING
By
Farhan Ahmad

A DISSERTATION

Submitted to
Michigan State University
in partial fulfillment of the requirements
for the degree of

DOCTOR OF PHILOSOPHY

Environmental Engineering

2011

ABSTRACT
RAPID DETECTION OF PATHOGENS AND THEIR ANTIBIOTIC SUSCEPTIBILITY
USING SIMPLE MICROFLUIDICS AND CCD IMAGING
By
Farhan Ahmad
Emergence of virulent and antibiotic resistant bacterial strains is posing a serious threat to
human health. In developing countries, deaths associated with infectious diseases (e.g., diarrheal
pathogens) are often due to the delayed diagnostics rather than effective and economical
treatment options. Gold standard methods for pathogen/resistance identification rely on either
polymerase chain reaction (PCR) or culture-based approaches, which are time consuming and
require either manual intervention(s) or expensive instrumentation. Microfluidic PCR-based
diagnostic systems require dedicated sample processing steps, thermal cycling, and sophisticated
microfabrication processes, thus are slow and expensive. Microfluidic loop-mediated isothermal
amplification (LAMP) is a promising alternative as it requires simpler instrumentation and
inexpensive detector due to its isothermal character and high amplicon yield. Therefore, for this
dissertation, novel fluorogenic dyes were applied to isothermal genetic and cellular assays for
rapid and real-time detection of pathogens and their antibiotic susceptibility by using simple
microfluidics and charge-coupled device (CCD) imaging.
A critical and quantitative review on miniaturized nucleic acid amplification systems was
performed for the selection of an appropriate molecular assay, low cost material for microchip,
detection system, and parameters that influence the gene amplification time. Real-time
fluorescence LAMP (microRTf-LAMP) assays of 12 virulence genes of 6 diarrheal pathogens
were performed in cyclic olefin copolymer microchips and monitored by a $1000 CCD-based

fluorescence imaging system. Application of a highly fluorogenic Syto-82 dye and CCD
exposure control to microRTf-LAMP assays, increased their signal-to-noise ratios by 10-fold and
reduced their threshold times to 10-50-fold, providing the single DNA copy level sensitivity. A
novel digital LAMP method to track the in situ bacterial growth by gene amplification directly
from gram-negative (Escherichia coli) and gram positive (Enterococcus faecalis) bacterial cells
was developed. Propidium monoazide-based LAMP assay was performed to confirm the gene
amplification was only from viable cells. Digital LAMP assays were rapid (20 min) and sensitive
to single cell. Cost reduction of 50-fold was achieved by using polyester material for digital
LAMP assays. To reduce the time of antibiotic susceptibility testing (AST), a microchip-based
dynamic method for E. coli growth monitoring in the presence of a membrane intercalating
fluorogenic dye (FM5-95) was developed. A variety of dye concentrations and CCD exposure
times were tested. A combination of 20 µg/mL FM5-95 dye and 10 s of CCD exposure provided
4

the detection of 1 ×10 E. coli cells in only 60 min. MicroAST of ampicillin and tetracycline on
E. coli strain provided the minimum inhibitory concentration (MIC) of 16 µg/mL and 1.6 µg/mL
respectively, which were consistent with the MIC values reported in literature.
The methods developed for this dissertation would eliminate the requirement of complex
sample preparation steps, facilitating the development of sensitive, rapid, low-cost, and
integrated diagnostic systems. The methods developed here for digital LAMP and microAST
when integrated on a single microfluidic chip will enable faster detection of AST in clinical
settings. The overall approach of faster detection of pathogens and their antibiotic resistance
based on pathogen growth in the presence of antibiotic(s) using digital LAMP directly on cells
can be extended to many different fields.

ACKNOWLEDGEMENTS

I feel surprised and the same time very grateful for everything I have received in my
professional life. Looking back at the past five years, they have been tough but shaped me as a
different person. The perspectives and skill sets, I have developed during this training period,
would certainly be useful for my future career and goals.
Any of my accomplishments during PhD would not have been possible without the help
and support of my advisor, Dr. Syed A. Hashsham. I believe that thanks would not be enough for
his mentorship. But still, I would like to thank him for having the confidence in me to do this
work. I still remember my first oral presentation in our group meeting and the critiques that I had
received from my advisor. It was due to his dedication and mentorship that helped me to grow up
as a scientist and have my own vision for research. I also truly appreciate his patience and
tolerance. He was always there whenever I needed him. I also thank my committee members Dr.
James Tiedje, Dr. Alison Cupples and Dr. Irene Xagoraraki for their valuable suggestions and
insights. I was fortunate enough to have them as my teachers. Their dedication and passion
towards teaching, research, and outreach services had always motivated me. Also, my special
thanks to other scientists and most importantly the best teachers, I have had at Michigan State
University, Dr. Terence Marsh, Dr. Volodymyr Tarabara, Dr. Thomas Voice, Dr. Phanikumar
Mantha, Dr. Jay Lennon, Dr. Edward Walker, and Dr. Russell Freed.
I would also like to thank for the all the support from Department of Civil and
Environmental Engineering and Graduate School at Michigan State University for the
fellowships and teaching assistantships during my PhD.

iv

It was a pleasure to share my doctoral studies and time with wonderful people like Bob,
Greg, Tiffany, Dieter, Prianca, Amanda, Maggie, Erick, Tim, and Aaron, who are very close
friends now. I would specially like to thank Bob for helping me in research, Dieter for his critical
comments, and Erick and Tim for useful discussions about antibiotic resistance and gene
sequencing. My days in East Lansing would not have been as enjoyable without some good
friends, including Rehan Baqri, Abu Hassan, Basir Ahmad, Pramod Thupaki, Kashi Telsang,
Indumathy Jayamani, and Weimin Sun. I would also like to acknowledge the help of the staff in
the CEE and CME office including Lori Larner, Margaret Conner, Laura Taylor, Yanlyang Pan,
Joseph Nguyen, and Maryam Thompson.
Finally, and importantly, I thank my family for their selfless support. I cannot imagine
going through this journey without the love and support of my wife Shazia, who has always
stood beside me in sharing my difficulties and celebrating my achievements.

v

TABLE OF CONTENTS

LIST OF TABLES ......................................................................................................................... ix
LIST OF FIGURES ........................................................................................................................ x
1. INTRODUCTION ...................................................................................................................... 1
1.1 Emergence of pathogens with antibiotic resistance .............................................................. 1
1.2 Gold standard methods for pathogen/resistance diagnosis ................................................... 2
1.3 Microfluidic-based molecular diagnostic tools ..................................................................... 3
1.4 CCD-based imaging .............................................................................................................. 6
1.5 Objectives ........................................................................................................................... 11
1.6 Organization ........................................................................................................................ 12
1.7 References ........................................................................................................................... 18
2. CRITICAL EVALUATION OF MINITURIZED NUCLEIC ACID AMPLIFICATION
SYSTEMS FOR POINT-OF-CARE DIAGNOSTICS ................................................................. 23
2.1 Introduction ......................................................................................................................... 23
2.2 Static chamber microPCR systems ..................................................................................... 28
2.2.1 External heaters ............................................................................................................ 29
2.2.2 Integrated heaters ......................................................................................................... 36
2.2.3 Non-contact heaters ..................................................................................................... 39
2.2.4 Analysis of static chamber microPCR systems ........................................................... 41
2.3 Flow-through microPCR systems ....................................................................................... 46
2.3.1 Analysis of flow-through microPCR systems.............................................................. 53
2.4 MicroPCR systems with innovative concepts..................................................................... 55
2.5 Surface treatment of microPCR systems ............................................................................ 58
2.6 Integrated microPCR systems ............................................................................................. 61
2.7 Isothermal nucleic acid amplification methods on microchip platforms ............................ 65
2.7.1 MicroLAMP systems ................................................................................................... 66
2.7.2 Instrument-free LAMP systems ................................................................................... 68
2.7.3 Simple sample processing in LAMP reactions ............................................................ 69
2.8 Conclusions ......................................................................................................................... 71
2.9 References .......................................................................................................................... 74
3. A CCD-BASED FLUORESCENCE IMAGING SYSTEM FOR REAL-TIME LOOPMEDIATED ISOTHERMAL AMPLIFICATION-BASED RAPID AND SENSITIVE
DETECTION OF WATERBORNE PATHOGENS ON MICROCHIPS .................................... 92
3.1 Introduction ......................................................................................................................... 92
3.2 Materials and methods ........................................................................................................ 95
3.2.1 DNA targets ................................................................................................................. 95
vi

3.2.2 LAMP primer design ................................................................................................... 95
3.2.3 Microchip fabrication................................................................................................... 96
3.2.4 Experimental setup....................................................................................................... 97
3.2.5 DNA dilution series and real-time LAMP ................................................................... 99
3.2.6 Fluorescence measurement, extraction, and analysis ................................................ 100
3.3 Results and discussion ...................................................................................................... 101
3.3.1 CCD optimization ...................................................................................................... 101
3.3.2 Effect of CCD exposure time on the fluorescence signals of dsDNA standard dilution
series ................................................................................................................................... 102
3.3.3 Effect of CCD exposure time on SNR and Tt values of microRTf-LAMP assays .... 105
3.3.4 MicroRTf-LAMP assays of waterborne pathogens ................................................... 110
3.3.5 MicroRTf-LAMP assay for quantitative analysis of C. jejuni ................................... 112
3.4 Conclusions ....................................................................................................................... 114
3.5 References ......................................................................................................................... 116
4. DIGITAL LAMP FOR RAPID AND SENSITIVE GENETIC AMPLIFICATION OF GRAMNEGATIVE AND GRAM-POSITIVE BACTERIA WITH NO SAMPLE PROCESSING ..... 120
4.1 Introduction ....................................................................................................................... 120
4.2 Materials and methods ...................................................................................................... 123
4.2.1 Bacterial cell and genomic DNA ............................................................................... 123
4.2.2 LAMP primer design ................................................................................................. 124
4.2.3 Microchip fabrication................................................................................................. 124
4.2.4 Experimental setup..................................................................................................... 125
4.2.5 RTf-LAMP on commercial PCR instrument and on microchip ................................ 126
4.2.6 PMA-RTf-LAMP assays with E. faecalis cells ......................................................... 127
4.2.7 Digital LAMP assays for most probable number calculation .................................... 128
4.2.8 Fluorescence signal analysis ...................................................................................... 128
4.3 Results and discussion ...................................................................................................... 129
4.3.1 RTf-LAMP assays from genomic DNA templates and bacterial cells ...................... 129
4.3.2 RTf-LAMP assays for quantitative analysis of E. coli and E. faecalis cells ............. 133
4.3.3 Quantification of viable E. faecalis cells by PMA-RTf-LAMP ................................ 136
4.3.4 MicroRTf-LAMP assays of single cell of E. coli and E. faecalis .............................. 140
4.3.5 Quantification of most probable number by digital LAMP ....................................... 143
4.4 Conclusions ....................................................................................................................... 147
4.5 References ......................................................................................................................... 149
5. RAPID AND SENSITIVE DETECTION OF THE GROWTH OF ESCHERICHIA COLI
AND ANTIBIOTIC SUSCEPTIBILITY TESTING ON MICROCHIP .................................... 155
5.1 Introduction ....................................................................................................................... 155
5.2 Methods and materials ...................................................................................................... 157
5.2.1 Microchip fabrication................................................................................................. 157
5.2.2 Fluorescence imaging system and signal acquisition ................................................ 157

vii

5.2.3 Bacterial cell and genomic DNA ............................................................................... 158
5.2.4 Bacterial growth and AST on microchip ................................................................... 158
5.2.6 Fluorescence signal analysis ...................................................................................... 160
5.3 Results and discussion ...................................................................................................... 160
5.3.1 Real-time fluorescence monitoring during E. coli growth......................................... 160
5.3.2 Effect of FM5-95 dye concentration and CCD exposure time on the Tt value of E. coli
growth on microchip ........................................................................................................... 164
5.3.3 Quantitative analysis of E. coli dilution series on microchip .................................... 165
5.3.4 Ampicillin and tetracycline susceptibility testing of E. coli in microchip ................. 168
5.3.5 Determination of MIC values of ampicillin and tetracycline for E. coli on microchip
............................................................................................................................................. 172
5.4 Conclusions ....................................................................................................................... 174
5.5 References ......................................................................................................................... 176
6. CONCLUSIONS AND FUTURE PERSPECTIVES ............................................................. 180
6.1 Conclusions ....................................................................................................................... 180
6.2 Future perspectives ........................................................................................................... 182
APPENDICES ............................................................................................................................ 185

viii

LIST OF TABLES

Table 1.1 Comparison of the parameters of monochrome CCD cameras .................................... 10
Table 2.1 Selected parameters of static chamber microPCR systems .......................................... 31
Table 2.2 Calculated thermal mass of thin film heaters for static chamber microPCR systems .. 45
Table 2.3 Parameters for systems with flow-through design including flow-through microPCR
systems .......................................................................................................................................... 50
Table 2.4 Examples of commonly used surface treatment strategies and their effect on microPCR
efficiency....................................................................................................................................... 60
Table 2.5 Selected examples of isothermal nucleic acid amplification methods ......................... 66
Table 3.1 Genes selected for detection of pathogenic microorganisms........................................ 96
5

Table 3.2 Comparison of average SNR and Tt values of RTf-LAMP for 10 DNA copies of V.
cholera toxR gene and C. parvum gp60 gene on microchips (1 s, 3 s, and 5 s CCD exposure
time) and the real-time PCR instrument. Standard deviations are the mean of Tt values from
triplicates ..................................................................................................................................... 106
5

Table 3.3 Comparison of average Tt values of RTf-LAMP for 10 DNA copies of 6 waterborne
pathogens (2 genes for each) on the microchips with 5 s CCD exposure and the real-time PCR
instrument. Standard deviations are the mean of Tt values from triplicates ............................... 111
Table A1.1 Nucleotide sequences of the designed LAMP primers ............................................ 185

ix

LIST OF FIGURES

Figure 2.1 Comparison of PCR time for static chamber microPCR chips with thin films (black
bars) and external heaters (shaded bars) ....................................................................................... 42
Figure 2.2 Effect of heating rates () on PCR times () for static chamber microPCR systems.
Inset shows the relation between thermal mass of thin film heaters and PCR time ..................... 44
Figure 2.3 Relationship between PCR time and the ratio of cross-sectional area and flow-rate.
Symbols and solid line corresponds, respectively to the data from studies reported on flowthrough microPCR systems and best fit line ................................................................................. 55
Figure 3.1 Photograph of the experimental setup consisting of an LED attached with excitation
filter (534 ± 20 nm) for illumination, a thin film heater and thermocouple for temperature
control, and a monochromatic CCD camera with emission filter (572 ± 20 nm) for imaging. The
inset shows the microchip with seven V-shaped reaction wells used for RTf-LAMP reactions.
This system was placed in the dark during imaging to avoid any ambient light .......................... 98
Figure 3.2 Optimization of offset and gain for the CCD camera. Plot of average fluorescence
signal from 10 ng dsDNA standard stained with 2 µM Syto-82 with respect to increasing offset
and gain at a fixed exposure time of 1 s...................................................................................... 102
Figure 3.3 CCD images of 16 well microchip with dsDNA standard dilution series (10 ng, 1 ng,
0.1 ng, 0.01 ng, and negative control) stained with 2 µM Syto-82 and with increasing exposure
times ............................................................................................................................................ 103
Figure 3.4 Plot of average fluorescence signal with respect to increasing exposure times for the
dilution series of dsDNA standard (10 ng, 1 ng, 0.1 ng, 0.01 ng, and 0 ng) stained with 2 µM
Syto-82 ........................................................................................................................................ 105
5

Figure 3.5 RTf-LAMP curves for 10 DNA copies of V. cholera toxR gene on the microchips (1
s, 3 s, and 5 s CCD exposure time) and real-time PCR instrument ............................................ 107
5

Figure 3.6 RTf-LAMP curves for 10 DNA copies of C. parvum gp60 gene on the microchips (1
s, 3 s, and 5 s CCD exposure time) and real-time PCR instrument ............................................ 107

x

Figure 3.7 Time lapse CCD images of microRTf-LAMP assay for V. cholera toxR gene with
CCD exposure times of 1 s, 3 s, and 5 s. Fluorescence signals increase with the increasing
exposure times ............................................................................................................................ 109
Figure 3.8 Standard curves for the C. jejuni 0414 gene amplification on the microchips at 5 s of
CCD exposure time and real-time PCR instrument .................................................................... 113
5

5

Figure 4.1 RTf-LAMP assays of 10 genomic DNA copies and 10 cells of E. coli (uidA gene)
o

and E. faecalis (gelE gene) at 63 C on a commercial PCR instrument ..................................... 131

Figure 4.2 SNR verses amplification time of the dilution series of RTf-LAMP assays of E. coli
5

(uidA gene) and E. faecalis (gelE gene) cells from 10 cells to single cell on a commercial PCR
instrument. For interpretation of the references to color in this and all other figures, the reader is
referred to the electronic version of this dissertation. ................................................................. 135

Figure 4.3 Standard curves of RTf-LAMP assays of E. coli (uidA gene) and E. faecalis (gelE
gene) cells on a commercial PCR instrument ............................................................................. 136
5

Figure 4.4 RTf-LAMP assays of 10 E. faecalis non-lysed cells (with and without PMA
treatment) and PMA treated lysed cells on a commercial PCR instrument ................................ 138
Figure 4.5 Colony forming units of non-lysed E. faecalis (with and without PMA treatment). No
growth was observed from heat treated cells .............................................................................. 139

Figure 4.6 A polyester chip (2.5 cm × 2 cm) for microRTf-LAMP assays. Each circular well was
1 mm in diameter with a volume of 2 µL. The circular wells were filled with food color ........ 140

Figure 4.7 MicroRTf-LAMP assays of single cell of E. coli (uidA gene) and E. faecalis (gelE
gene). Inset showed the CCD time lapse images of microchip with 10 positive and 10 negative
controls. The error bars represented the standard deviations of the mean from 5 and 8 positive
replicates and 10 negative replicates for E. coli and E. faecalis respectively ............................ 141

Figure 4.8 Single cell RTf-LAMP assay of E. coli (gadA gene) on a commercial PCR instrument.
Only two positive controls showed amplification out of six replicates. The error bars represented
the standard deviations of the mean from two replicate experiments ......................................... 144

xi

Figure 4.9 Digital LAMP image for E. coli (gadA gene) dilutions. Cell number per well (left)
and number of positive reaction (right) were shown in the image ............................................. 146
Figure 4.10 Digital LAMP image for E. faecalis (gel E gene) dilutions. Cell number per well
(left) and number of positive reaction (right) were shown in the image .................................... 146
Figure 5.1 Real-time fluorescence growth curves of E. coli with various concentrations of FM595 dye at 1 s and 10 s of CCD exposure times. .......................................................................... 162
4

Figure 5.2 Comparison of E. coli (1 ×10 cells) growth detection times (SNR cut-off ≥ 20) with
various concentrations of FM5-95 dye measured with 1 s and 10 s of CCD exposure times. ... 165
4

Figure 5.3 Real-time fluorescence growth curves of the dilution series of E. coli (1 ×10 cells to
1 cell) with 20 µM FM5-95 dye and the standard curve. For single cell dilution, the error bars
represent the standard deviations of the mean only from 2 replicates ........................................ 167
Figure 5.4 Chemical structures of (a) ampicillin and (b) tetracycline ........................................ 169
4

Figure 5.5 Real-time fluorescence growth curves of 1 ×10 cells of E. coli (with 20 µg/mL FM595 dye) and increasing concentrations of (a) ampicillin and (b) tetracycline ............................. 171
4

Figure 5.6 Ratio of the growth rates of 1 ×10 cells of E. coli (with 20 µg/mL FM5-95 dye) with
and without (a) ampicillin and (b) tetracycline ........................................................................... 173
Figure 6.1 Conceptual picture a microfluidic chip for integrated AST and digital LAMP ........ 182

xii

CHAPTER I
1. INTRODUCTION

1.1 Emergence of pathogens with antibiotic resistance
Emergence of virulent and antibiotic resistant (AR) bacterial strains is posing a serious
threat to human health (Soto 2009; Allen et al. 2010). According to the National Institute of
Allergy and Infectious Diseases, the annual treatment costs of AR infections in the U.S are $5
billon. The death rates by some of the major drug-resistant pathogens like tetracycline and
methicillin-resistant

Staphylococcus

aureus

(MRSA),

penicillin-resistant

Streptococcus

pneumoniae, vancomycin-resistant Enterococcus spp., and multidrug-resistant Mycobacterium
tuberculosis have risen drastically. Due to the consistent and excessive use of antibiotics in
clinical, agricultural, and livestock settings, targeted bacterial strains have evolved novel genetic
mechanisms to combat the actions of antibiotics (Nikaido 2009). Once acquired, antibiotic
resistance genes (ARGs) can be transferred vertically through host cell division and horizontally
to other bacterial species by transduction, conjugation, or transformation. These AR traits are not
only been transferred to bacterial species in local environment but also crossing global
boundaries. An example is the recent discovery of carbapenem-resistance gene (blaNDM-1) in
Klebsiella pneumoniae and Escherichia. coli strains which were imported into the UK and
Kenya by patients returning from the Indian subcontinent (Kumarasamy et al. 2010; Poirel et al.
2011). Another major problem is diarrheal diseases, which are the second leading cause of death,
killing approximately 2.2 million people worldwide, according to World Health Organization
(Craun et al. 2010a). Most of the cases of diarrheal diseases are confined to underdeveloped
nations due to contamination of food/water, lack of improper sanitation and late treatment, and

1

poverty. Approximately 60% of diarrheal diseases are caused by pathogenic bacteria and
protozoan including, Campylobacter, Salmonella, Shigella, E. coli O157:H7, Giardia, and
Cryptosporidium. Emergence of ARGs in these pathogens is posing additional challenges related
to their treatment (Sack et al. 1997). In many cases, deaths due to infectious diseases are because
of the lack of appropriate diagnosis rather than availability of effective and economical
prevention and treatment options (Yager et al. 2008). A recent estimation showed that a 100%
accurate diagnostic test for tuberculosis could save at least 625,000 lives per year globally
(McNerney and Daley 2011).
1.2 Gold standard methods for pathogen/resistance diagnosis
Gold standard methods for antibiotic susceptibility testing (AST) and pathogen
identification rely upon culture-based approaches, e.g., broth or agar dilution, and disk diffusion,
which are simple and inexpensive (Bauer et al. 1966; Stalons and Thornsberry 1975). However,
they are also selective and time consuming to permit same-day diagnosis and this delays
treatment and increase the risk of inappropriate use of antibiotics. Cell growth in conventional
culture methods has longer lag-phase (6 h for E. coli by absorbance measurement) and is often
limited by the depletion of nutrients and accumulation of metabolites (Groisman et al. 2005;
Arora et al. 2009). Towards the goal of achieving faster AST there is a need for improvement in
optical methods and development of microchip-based platforms to detect cell growth (Boedicker
et al. 2008; Chen et al. 2010). Relatively faster genetic methods such as polymerase chain
reaction (PCR) and loop-mediated isothermal amplification (LAMP) have been extensively
applied for the detection of pathogenic microorganisms and genetic markers for resistance
(Seyrig et al. 2010). However, these methods only provide the genetic profiles of the pathogens
and lack the ability to confirm the phenotypic functions such as susceptibility to antibiotics.
2

Additionally, AR trait will only be present upon the expression of ARG(s) in the AR pathogen.
Commercial instruments for AST and nucleic acid amplification/detection are highly expensive.
For example, turbidity-based VITEK® 2 (Biomerieux) and fluorescence-based autoSCAN® 4
(Siemens) for AST have detection time of 24 h and costs around $70,000. Commercial PCR
instruments like RAPID® System (Idaho Technologies) and RAZOR EX® Instrument (Smith
Detection) takes 1-2 h and costs in the range from $30,000 to $50,000. Also the cost per assay
range from $10 to $50 or higher depending on the reaction chemistry. The need for expensive
instrumentation and highly trained personnel, however, currently limits their widespread use,
especially in resource-poor settings.
1.3 Microfluidic-based molecular diagnostic tools
In resource-limited settings, molecular tests are commonly performed in centralized
laboratories. Applications where a point-of-care (POC) genetic analysis system will provide
superior capabilities compared to centralized screening are numerous. Examples include
quantitative detection of viral load for human immune deficiency syndrome (Richman et al.
2009), genotyping of influenza (Van Kerkhove et al. 2011), M. tuberculosis with capabilities to
detect drug resistance (McNerney and Daley 2011), MRSA (Klevens et al. 2007), malaria
(Plasmodium spp.) (Feachem et al. 2010), dengue (Guzman et al. 2010), genotyping for E. coli
O157:H7 (Craun et al. 2010b). Many of these applications also require screening for multiple
agents simultaneously. In resource-limited settings, immunoassay-based diagnostic tests for
these infectious agents are commonly performed due to their speed, simplicity, and low-cost,
albeit having lower sensitivity than culture methods and nucleic acid-based assays. For
molecular assays to become more easily accessible outside of laboratory settings, detection time,
complexity of the test, instrumentation cost, and cost-per assay must be reduced.
3

Early on, semiconductor manufacturing technology played a significant role in the
development of microchips made of silicon or glass for nucleic acid amplification. However,
because silicon wafers need for sophisticated fabrication facility, high turn-around time, and are
expensive, more recent efforts have shifted to a variety of polymeric materials (Fiorini and Chiu
2005; Bhattacharyya and Klapperich 2006; Tsao and DeVoe 2009; Ahmad et al. 2011;
Chantiwas et al. 2011). Application of polymers in DNA amplification microchips drastically
reduces the turnaround time from days to hours and cost of disposable microchips from dollars to
cents. Towards the development of miniaturized nucleic acid analysis systems for the automation
of steps from sample preparation through amplification reaction to result, microchips are being
integrated with additional components such as pumps, valves, and detectors (Zhang et al. 2007;
Liu et al. 2010). Partially integrated nucleic acid analysis systems with amplification and
detection modules, based on PCR and isothermal nucleic acid amplification assays (e.g., loopmediated isothermal amplification, LAMP) with nucleic acids as the starting material, has been
extensively reported (Zhang et al. 2006; Zhang and Da 2010; Asiello and Baeumner 2011).
However, seamless integration of all the analytical steps, including the onchip sample processing
from pathogenic bacterial/viral cells (e.g., cell lysis, DNA extraction and purification) is only
beginning to emerge (Chen et al. 2007; Park et al. 2011). Complete integration of all the
analytical processes in miniaturized nucleic acid analysis systems is must for a successful POC
genetic analysis device disease diagnostics (Lee et al. 2010; Chin et al. 2011; Niemz et al. 2011).
MicroPCR systems required dedicated nucleic acid processing modules due to possibility of
reaction inhibition by a variety of components present in raw samples. Thermal cycling time of
microPCR system is dependent on thermal mass of the system, which if reduced could lead to
faster gene amplification. For example, nucleic acid amplification on microPCR systems could

4

be performed from 90 s for a flow-through system (Kopp et al. 1998) to 6 min for a miniaturized
silicon integrated heater (Neuzil et al. 2006). On the other hand microLAMP system uses simpler
instrumentation due to its isothermal character and inexpensive detector due to high amplicon
yield (Seyrig et al. 2010; Ahmad et al. 2011). Translation of these characteristics on microchipbased platforms is expected to lead to low-cost miniaturized molecular devices for rapid
pathogen diagnosis. Development of molecular diagnostic devices for POC settings requires an
approach with a goal towards speed, sensitivity, specificity, user-friendliness, robustness, and
low cost.
Monitoring of nucleic acid amplification assays (e.g., PCR or isothermal variants) in
microchips are commonly performed by optical detection strategies, including absorbance,
fluorescence, interferometry, chemiluminescence, and surface plasmon resonance (Kuswandi et
al. 2007; Myers and Lee 2008). Fluorescence sensing is most common due to its high sensitivity
and selectivity as well as the commercial availability of well-established fluorophores (Schulze
et al. 2009) and because fluorescence forms the cornerstone for many molecular diagnostic
assays (Niemz et al. 2011). Sensitivity of optical devices is influenced by multiple parameters,
including intensity of light source, fluorescence light collection efficiency, and background
signal and noise. Fluorescence light collection efficiency is dependent on the light channeling
components (optical fibers, filters), light collection components (lens, detector) and their
assembly in the device. Inexpensive, compact and robust optical devices require low-cost, small
and durable light sources. Light emitting diodes (LEDs) meet these criteria and are very well
suited for diagnostic devices (Moe et al. 2005). LEDs are increasingly being used to replace
conventional light sources such as lasers, and mercury and tungsten-halogen lamps. Important
requirements from photon detectors are high sensitivity and signal-to-noise ratio (SNR), low

5

cost, and compactness. Photon detectors such as solid state photodiodes, vacuum-based
photomultiplier tubes, charge-coupled devices (CCDs) are used in conventional and microchipbased optical systems (Yotter and Wilson 2003). For portable microchip-based devices,
integration of intense light sources and sensitive photon detectors is critical, especially
considering the minute sample volumes. This is because reduced sample volumes in the microto nanoliter range also translate into smaller detection volumes and a decrease in the number of
analytes available for detection. The challenge in developing microchip-based detection systems
is the miniaturization of the optical hardware at low cost while still maintaining the sensitivity
and specificity required for interrogating molecular assays. Drastic improvements in these
systems can be realized with the improvement in the quality of light sources and photon
detectors and the availability of low-cost optical components. Leveraging these off-the-shelf
technologies may be one of the keys to realizing POC diagnostic devices.
1.4 CCD-based imaging
Compared to spot detectors such as photodiode and photomultiplier tube, CCD has wider
field-of-view, thus suitable for high throughput fluorescence imaging applications (Salmon and
Waters 2011). Also due to exposure control features in CCDs, they provide more sensitive
detection than simple photodiodes. CCD camera was invented at AT&T Bell Labs by Willard
Boyle and George E. Smith in 1969 (Boyle and Smith 1970). A CCD is an array of light-sensing
elements (pixels) arranged on a silicon substrate. The semiconductor properties of silicon allow
the light sensor to trap and hold photon-induced charge carriers (photoelectrons) under
appropriate electrical bias conditions. The basic light-sensing unit of the CCD is a photodiode
attached with a storage area for photoelectrons (potential well). Photodiode elements of the CCD
absorb the energy from incoming photons resulting into liberation of photoelectrons, and the
6

formation of holes within the silicon crystal lattice. A single electron-hole pair is generated from
each absorbed photon. The numbers of photoelectrons generated at a particular pixel is directly
proportional to the incoming light intensity. Each pixel in the sensor accumulates these charges
during the imaging process. During readout, these collected charges are subsequently shifted
along the transfer channels under the influence of applied voltages. Finally, an analog to digital
convertor assigns a digital value to each pixel according to its voltage amplitude, creating a
digital image.
Characteristics of CCD camera such as sensor architecture, spatial resolution, dynamic
range, quantum efficiency, exposure time, and noise factor are the critical determinants of the
quality of digital image and the resulting SNR, and detection sensitivity (Hiraoka et al. 1987;
Salmon and Waters 2011). Spatial resolution of a CCD sensor is a function of the number of
pixels and their size. Spatial resolution is also affected by the quality of the lens and imaging
optics. Readout noise is associated with individual pixel. The total readout noise will be
dependent on the total number of pixels on the CCD sensor. Therefore, the readout noise will be
higher for relatively high numbers of smaller-sized pixels than low numbers of relatively larger
pixels. A proper balance is needed between the required resolution and the SNR. Binning is
commonly performed to increase the SNR by compromising with the spatial resolution. Binning
is the process of combining pixels to form a super pixel. A binning of 2×2 means that an area of
4 adjacent pixels is combined into one larger pixel. In this case, the SNR will increase by 4 times
and the resolution of the image will be reduced by half. CCDs are available with a range of pixel
sizes (1.4 µm-30 µm) and pixel numbers (1-50 megapixels). For example, Kodak recently
launched a full frame CCD sensor (KAF-50100) with highest number of pixels (50 megapixels)
with unit pixel dimension of 6 µm × 6 µm. Sony has a 14 megapixel interline CCD sensor

7

(ICX681SQW) with smallest pixel dimension of 1.43 µm × 1.43 µm. Dynamic range of a CCD
is a measure of two properties; i) the difference between the highest and lowest signal values,
and ii) the number of steps captured between the highest and lowest values. A CCD sensor with
larger dynamic range can capture faint signal without distorting its highlights in the image.
Dynamic range is calculated by dividing the full well electron capacity by the readout noise. Full
well electron capacity is number of total photoelectrons a pixel can hold before saturation. For
example, a CCD with full well electron capacity of 55,000 per pixel and readout noise of 8 ewill have a dynamic range of 6875. The signals lower than the readout noise (<5 e-) will be
buried in the noise, thus readout noise represents the smallest possible signal this CCD can
capture. Dynamic range also represents the number of steps the CCD is capable of recording
above the noise. Bit depth, which is the property of CCD’s analog to digital converter,
transforms those captured steps into digital format. Conversion of all the captured steps in the
digital format allows the image to have clear signal. For example, a CCD with 12 bit depth and
16 bit depth will have 2

12

(4,096) and 2

16

(65,536) digital steps respectively. CCD with higher

bit depth will capture more variations or shades in the image. CCD with 12 bit depth (4,096
digital steps) will not utilize the available dynamic range of 6875. However, a CCD with 16 bit
depth (65,536 digital steps) will take full advantage of the available dynamic range. Therefore,
CCD with high dynamic range is needed to capture fine gradations of intensity steps in the
image. At the same time, a high bit depth is needed to convert all of those captured steps into a
usable digital output. Quantum efficiency is the ability of the CCD to convert incoming photons
into signal output. It is defined as the ratio of incoming photons to those actually detected in the
device. Silicon has band gap energy of 1.14 electron volts. It can absorb the photons with
wavelengths ranging from 400 nm-1100 nm. Silicon will be transparent to the wavelengths lower
8

than 400 nm (absorption by gate structures) and higher than 1100 nm (low energy for electronhole generation). Quantum efficiency is measured over a range of different wavelengths to
characterize the CCD efficiency at different photon energies. Quantum efficiency of
commercially available CCDs lies between 40-50%. Exposure/integration time is directly related
to the measured signal from a CCD. Scientific CCDs have exposure times from milliseconds to
hours. Emission light from low sample volumes is also very low (100-1000 photons). In order to
image the low signals, a high exposure is needed. It has been demonstrated that increasing the
exposure time increases the SNR. However, increasing the exposure time will also increase the
background signal and dark current along with the fluorescence signal. It is suggested to cool the
CCD sensor for exposure time higher than 1 second. Cooled CCDs have successfully been
applied to detect microchip-based chemiluminescence assays with an exposure time of 20 min
(Yang et al. 2008, 2009). Noise factors are introduced during the integration and read out of an
image from a CCD. Three major noise factors are shot noise, read noise and dark current
(Robbins and Hadwen 2003). CCD manufactures report these noise factors as electrons root
mean square or number of electrons per pixel. Shot noise is inherent to light characteristics and
results from the statistical variation in the arrival rate of photons on CCD. The probability of
photon arrival on a pixel is predicted by Poisson distribution. Shot noise is directly proportional
to the square root of signal value. It is more apparent for low light conditions. It can be reduced
by either collecting more photons (increase in exposure time) or by combining multiple frames.
Readout noise is a form of electronic noise added to the final signal upon readout of the device,
is independent of exposure time. Read out noise can be suppressed by binning the pixels. Dark
current is a multiplicative form of noise, which is proportional to the length of the exposure time.
It is caused by thermally generated electrons that build up in the pixels and depends on the

9

exposure time and temperature of the CCD sensor. It is the square root of the number of thermal
electrons multiplied by the exposure time. It has been reported that for cooling the CCD for
every 6°C will reduce the dark current by half. MEADE DSI Pro II applies a unique concept of
dark frame subtraction to reduce the dark current. Dark frames are the images taken in the dark
with different exposure times. During imaging, this dark frame is automatically subtracted from
an image by matching the imaging exposure time with the exposure time of the stored dark
frame.

Table 1.1 Comparison of the parameters of monochrome CCD cameras
Parameter
Company

Non-cooled CCD
TM
Meade
DSI Pro
Sony ICX429ALL
Interline
752 × 582
8.3 × 8.6
50,000
16
0.1 - 1
15
none

DSI Pro (MEADE

EMCCD
TM
Proxitronic
Proxi eCam 200
TI TC247SPD
Interline
658 × 496
10 ×10
14
0.5 - 0.04
≤1
-20

283
399

Model
Sensor
CCD type
Pixels
Pixel size (µm)
Full well capacity (e/p)
ADC (bit)
Exposure time (ms-h)
Read noise (e)
o
Cooling ( C)
Weight (g)
EM gain
Cost ($)

Cooled CCD
TM
ATIK
16 IC-HS
Sony ICX424AL
Interline
659 × 494
7.4 ×7.4
30,000
16
1-1
7
-20
350
495

500
2000
11,000

TM

TM

) and 16 IC-HS (ATIK

) cameras were is very similar in terms of

the given specifications expect the peltier cooling in 16 IC-HS, which reduced the noise factors
to half but increased the cost by $100 (Table 1.1). Another difference was the better dynamic
range (4286) of 16 IC-HS than DSI Pro (3333). Electron multiplying CCD (Proxi eCam 200)
was 25-times expensive than DSI Pro and 16 IC-HS. DSI Pro and 16 IC-HS were applied to

10

image PCR and chemiluminescence assays in microfluidic chips (Kaigala et al. 2008; Yang et al.
2008, 2009). Also, the effect of CCD exposure time on increasing SNR was demonstrated.
However, these low-cost CCDs were not yet applied for monitoring microRTf-LAMP assays.
Therefore, we selected DSI Pro CCD to evaluate the effect of exposure time on SNR and
threshold time (Tt) of microRTf-LAMP assays (Ahmad et al. 2011).

1.5 Objectives
The overall goal of this dissertation research was to develop a simple platform for rapid
detection of pathogens and their antibiotic susceptibility. The specific objectives are discussed
below;

Objective 1: To reduce Tt of microRTf-LAMP by 40% as compared to real-time PCR
instrument by using a novel DNA intercalating dye and a low-cost CCD imaging system.
Hypothesis: Application of CCD exposure control to Syto-82-based microRTf-LAMP may
reduce Tt due to increase in the SNR by high fluorescence of Syto-82 dye and with increasing
CCD exposure time (as already been shown for PCR and chemiluminescence assays).

Objective 2: To develop digital LAMP method for amplifying genetic sequences directly from
o

bacterial cells without any processing at a moderate temperature of 63 C.
o

Increase in the temperature about 50 C slightly increases the membrane permeability of gramnegative and gram-positive bacterial cells.

11

Hypothesis: In situ LAMP may be performed, considering the successful inclusion of LAMP
reaction components inside the bacterial cell even with a slight change in membrane
o

permeability at 63 C.
Objective 3: To develop a dynamic method of measuring cell growth/ antibiotic susceptibility
and reduce the Tt to 1-2 h by using a novel cell membrane intercalating dye and low-cost CCD
imaging system.

Hypothesis: Application of CCD exposure control to FM5-95-based microAST may reduce Tt
due to increase in the SNR by high fluorescence of FM5-95 dye and with increasing CCD
exposure time.
1.6 Organization
This dissertation is organized into six chapters focusing on various aspects of the overall
project. Chapter I introduces the needs and requirements for rapid, low-cost, and simplified
miniaturized systems for the diagnosis of pathogens and their antibiotic susceptibility. Chapter II
presents an extensive critical review on the miniaturized molecular diagnostic systems such as
microPCR and microLAMP. The main aim of this review is to understand the effect of various
parameters (e.g., thermal mass and conductivity, heating/cooling rate) of the nucleic acid
amplification time and the current state of these systems in terms of their advantages and
limitations for POC diagnostics. This quantitative comparison of miniaturized molecular
diagnostic systems is used for the selection of an appropriate nucleic acid amplification assay,
microchip material, and detection system for this project. Chapter III (published) presents the
development of Syto-82 dye-based LAMP assays for 12 virulence genes of 6 diarrheal
pathogens, translation of these assays on microfluidic chips, and evaluation of CCD exposure
12

time on the SNR and threshold time of microLAMP assays. Chapter IV presents the
development of novel microLAMP assays, “digital LAMP”, for the genetic amplification from
o

E. coli (uidA gene) and E. faecalis (gelE gene) cells at 63 C without any sample processing.
Chapter V evaluates a membrane intercalating dye, FM-5-95 and CCD exposure time, for the
rapid detection of E. coli growth on microchip. The concept is also extended for the rapid testing
of antibiotic sensitivity (ampicillin and tetracycline) on E. coli. Finally, Chapter VI presents the
conclusions and future directions. The designed LAMP primer sequences are provided in the
appendix at the end. Brief summary of these chapters is provided below.

Chapter II Critical evaluation of miniaturized nucleic acid amplification systems for POC
diagnostics: POC genetic diagnostics critically depends on miniaturization of amplification,
detection, and sequencing techniques. Isothermal amplification techniques, which are still
emerging, have a better potential for POC diagnostics. Isothermal amplification techniques (e.g.,
LAMP) are promising for the low-cost and integrated nucleic acid analysis due to i) moderate
incubation temperature leading to simplified heating and low power consumption, ii) high
amount of amplification products, which can be detected visually or by simple detectors, iii)
direct genetic amplification from bacterial cells due to the superior tolerance to substances that
typically inhibit PCR, iv) high specificity, sensitivity up to single copy, and rapid detection in
10-20 min. LAMP assays on microfluidic chips (microLAMP) were evaluated under the World
Health Organization’s guidelines of POC diagnostic systems towards elimination of any
instrumentation and integration of sample processing step in future. Microfluidic chips for
microPCR with different materials and designs have been reported. Temperature cycling systems
with varying thermal masses and conductivities, thermal cycling times, flow-rates, and crosssectional areas, have also been developed to reduce the nucleic acid amplification time. A
13

quantitative critical evaluation of these factors in relation to time taken to amplify nucleic acid
was not available in the current literature. An in-depth analysis of the parameters responsible for
rapid nucleic acid amplification in two major types of microPCR systems (static chamber and
flow-through) along with their integration with sample processing steps for sample-in-answerout capability was presented. The major aim was to provide a better understanding of the
advantages and limitations of microPCR and microLAMP systems for POC diagnostics.

Chapter III A CCD-based fluorescence imaging system for real-time loop-mediated
isothermal amplification-based rapid and sensitive detection of waterborne pathogens on
microchips: Rapid, sensitive, and low-cost pathogen diagnostic systems are needed for early
disease diagnosis and treatment, especially in resource-limited settings. This study reports a lowcost CCD-based imaging system for rapid detection of waterborne pathogens by isothermal gene
amplification in disposable microchips. Fluorescence imaging capability of this monochromatic
CCD camera was evaluated by optimizing the gain, offset, and exposure time. This imaging
system was validated for 12 virulence genes of major waterborne pathogens on cyclic olefin
copolymer microchips, using Syto-82 dye and real-time fluorescence LAMP referred here as
microRTf-LAMP. SNR and Tt values of microRTf-LAMP assays were compared with those
from a commercial real-time polymerase chain reaction (PCR) instrument. Applying a CCD
5

exposure of 5 s to 10 starting DNA copies of microRTf-LAMP assays increased the SNR by 8fold and reduced the Tt by 9.8 min in comparison to a commercial real-time PCR instrument.
Additionally, single copy level sensitivity for Campylobacter jejuni 0414 gene was obtained for
microRTf-LAMP with a Tt of 19 min, which was half the time of the real-time PCR instrument.
Due to the control over the exposure time and the wide field imaging capability of CCD, this
14

low-cost fluorescence imaging system has the potential for rapid and parallel detection of
pathogenic microorganisms in high throughput microfluidic chips.

Chapter IV Digital LAMP for rapid and sensitive genetic amplification of gram-negative
and gram-positive bacteria with no sample processing: An isothermal method for gene
amplification directly from gram-negative (Escherichia coli) and gram positive (Enterococcus
o

faecalis) bacterial cells at 63 C with most probable number (MPN) approach leading to the
5

concept of “digital LAMP” is reported. RTf-LAMP assays (from 10 cells to single cell) of E.
coli (uidA gene) and E. faecalis (gelE gene) were successfully performed by directly incubating
0

the cells at 63 C (without any prior processing) on a commercial PCR instrument. Threshold
o

time values of RTf-LAMP assays of bacterial cells, heat treated bacterial cells (95 C for 5 min),
and their purified genomic DNA templates were similar, implying that DNA could be amplified
o

either by lysis or permeabilization of bacterial cells at 63 C. Gene amplification from viable E.
faecalis cells was confirmed by performing propidium monoazide (PMA)-RTf-LAMP assays.
Single cell-based RTf-LAMP assays were also translated on polymeric microchips (microRTfLAMP) and monitored by a low-cost fluorescence imaging system. MicroRTf-LAMP assays
provided SNR ratio between 500- to 800, sensitivity of single cell, and detection time of
5

approximately 20 min. Endpoint microLAMP assays were performed from 10 bacterial cells to
0.01 bacterial cell in digital format for MPN analysis. MPN values from digital LAMP assays
confirmed the genetic amplification were from single cell. Direct gene amplification from
o

bacterial cells at 63 C would eliminate the requirement of complex sample preparation steps,
15

facilitating the development of sensitive, rapid, low-cost, and integrated POC diagnostic systems.
In addition, rapid digital LAMP has the potential to replace conventional MPN method.

Chapter V Rapid and sensitive detection of the growth of Escherichia coli and antibiotic
0

susceptibility testing on microchip: Real-time growth monitoring of E. coli at 37 C in the
presence of a membrane intercalating FM dye, by using a low-cost microchip and CCD imaging
system is reported. Various concentrations of FM5-95 dye (1 µg/mL-50 µg/mL) and two
different CCD exposure times (1 s and 10 s) were tested towards the rapid detection of E. coli
growth. A combination of 20 µg/mL FM5-95 dye and 10 s of CCD exposure time provided the
4

detection of 10 E. coli in only 1 h. Highest concentration of FM5-95 (50 µg/mL) inhibited the
E. coli growth. Sensitivity of the cell growth assay was tested with the dilution series of E. coli
4

from 10 to 1 cell. The assay was sensitive to single cell with a detection time of 5 h. Antibiotic
sensitivity of ampicillin and tetracycline on this E coli strain was also tested. MIC values of
ampicillin and tetracycline were 16 µg/mL and 1.6 µg/mL respectively, which were also
consistent with the literature values. This method would be useful for high throughput, rapid and
real-time monitoring of bacterial growth for their AST.

16

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22

CHAPTER II
2. CRITICAL EVALUATION OF MINITURIZED NUCLEIC ACID AMPLIFICATION
SYSTEMS FOR POINT-OF-CARE DIAGNOSTICS
2.1 Introduction
Miniaturized nucleic acid amplification systems are highly needed for point-of-care
(POC) diagnostic applications. For genetic testing, two key approaches, that are commonly used
are polymerase chain reaction (PCR) and isothermal assay (e.g., loop-mediated isothermal
amplification, LAMP) (Notomi et al. 2000b). Other approaches that are potentially useful but
have received less attention include transcription-mediated amplification (Hofmann et al. 2005),
helicase dependent amplification (Vincent et al. 2004), recombinase polymerase amplification
(Piepenburg et al. 2006), nucleic acid sequence-based amplification (Gracias and McKillip
2007), and non-amplification-based method (Nam et al. 2004). Early studies in this direction
started in 1993 with the demonstration of PCR in 50 µL silicon microchip (Northrup et al.
1993b). Over the last two decades, there has been a surge in miniaturized nucleic acid
amplification systems with innovative concepts, designs, sizes, materials, modes of integration
among others (Zhang et al. 2006; Zhang and Xing 2007a; Zhang and Ozdemir 2009; Park et al.
2011). Molecular assays carried out in micro/microfluidic chips typically enable shorter analysis
times, reduce reagent consumption, minimize risk of sample contamination, and often enhance
assay performance (except lower limit of detection due to smaller sample volume).
MicroPCR/LAMP refers to microfluidic chip with micron size chamber(s) for performing single
or multiple PCR or LAMP reactions (Mark et al. 2010). MicroPCR/LAMP system in its most
basic form includes a microchip for confined reaction and a heater for thermal cycling. In

23

addition, it may also include pumps and valves for liquid handling, detection components, and
interface for heater control and other parameters (Burns et al. 1998). PCR is an enzymatic
process of amplifying selected regions of nucleic acid (Saiki et al. 1988). It relies on thermal
cycling i.e., cycles of repeated heating and cooling of the PCR reagent mixture (target DNA,
nucleotides, polymerase, primers, and buffer components). This process produces large number
of copies of a specific segment of target DNA. LAMP is a promising alternative since the
o

amplification reaction is performed at 63 C with the strand displacement activity of polymerase
(Notomi et al. 2000b). LAMP provides higher amplicon yield than PCR, thus the positivity of
amplification can be detected either visually or by low-cost detectors. MicroPCR/LAMP systems
have the potential of rapid amplification, multiplexing, portability, and low-cost integration for
complete nucleic acid analysis.
Static chamber and flow-through microPCR systems are commonly applied for the
amplification and detection of nucleic acids. Based on whether the temperature is cycled for PCR
reagent mixture contained in a static chamber or it is cycled through various temperature zones,
microPCR systems can be classified into two major types, static chamber type (Northrup et al.
1993a) and flow-through type (Kopp et al. 1998). Other types of microPCR systems such as
convective, thermal gradient, and oscillatory are also reported. Static chamber microPCR chip is
commonly fabricated in well or capillary format. Flow-through microPCR chip is fabricated in a
serpentine channel format to take advantage of the three temperature zones. Certain parameters
in microPCR systems like thermal mass, conductivity, flow rate, and channel cross-sectional
area, which if optimized properly can lead to amplification in less than 5-10 min compared to 30
min to 2 hours in conventional PCR systems. Using the flow-through system, it can be reduced
8

to as low as 1.5 min, albeit for large starting copy numbers (10 copies) (Kopp et al. 1998).
24

There are a number of advantages of static chamber microPCR over the conventional PCR like
rapid gene amplification, lower cost of reagents and system components, high throughput, and
possibility of integration as a portable or hand-held molecular diagnostic device (Zou et al. 2002;
Lee and Hsing 2006; Wang et al. 2006). Comparatively, more reports are available on static
chamber microPCR systems than flow-through microPCR systems. More interest of the research
community towards static chamber microPCR systems is due to: i) smaller footprint of chip, ii)
possibility of high throughput, and iii) easier integration of detection step for real-time nucleic
acid amplification (Xiang et al. 2005b). The concept of real-time flow-through microPCR system
has been often described in literature but only one real-time flow-through microPCR system has
been reported yet (Pjescic et al. 2010a).
Static chamber microPCR systems have similar working principle as commercial PCR
machines in which 50-100 l vials containing PCR reagent mixture are placed on a metallic
o

block that can undergo through a temperature cycle causing target DNA denaturation (95 C for
o

o

1 min), primer annealing (55 C for 1 min), and copying or extension (72 C for 2 min). This
cycle is repeated 30 to 35 times, which takes around 2 h for complete nucleic acid amplification.
During this process, time is required for bringing the temperature of the heating block to the
cycle equilibrium temperature and to transfer heat to the PCR reagent mixture. Due to the high
mass and limited response time of the heating blocks in commercial PCR machines, until
o

recently very slow heating and cooling rates (1-2 C /s) were employed. This, however, is
o

changing with systems that now have 10-20 C /s of heating and cooling rates requiring only 2030 minutes for nucleic acid amplification. In both the cases, a major fraction (>90%) of the total
time is needed due to ramping up and down of the heating block rather than the amplification
25

process. It has been shown that DNA denaturation and primer annealing are almost instantaneous
(< 1 s) once the PCR reagent mixture reaches to appropriate temperature (Saiki et al. 1988).
Additionally, Taq polymerase commonly used in PCR has a nucleotide extension rate of ~100
nucleotides/s (Pavlov et al. 2004). Therefore, considering 1 sec each for denaturation, annealing,
and extension, it is possible to amplify a 100 bp amplicon in 90 sec provided the time for
temperature cycling can be eliminated. A number of DNA polymerases with much higher
enzymatic activity are also known today and are becoming the basis for rapid PCR (Pavlov et al.
2004). Nucleic acid amplification in commercial PCR systems is characterized by high thermal
mass, low temperature ramping rates, high amplification time, and high power consumption
(Wittwer et al. 1989a). In contrast to commercial PCR machine, static-chamber microPCR has
smaller sample volume (nL to µL) and smaller foot print of microchip and heater. Flow-through
microPCR does not require temperature cycling, resulting in a much faster amplification than
static chamber microPCR system. In flow-through microPCR, the PCR reagent mixture flows
o

o

o

through a channel passing through the three temperature zones (95 C, 55 C, and 72 C).
During the flow, the PCR reagent mixture comes in thermal equilibrium with the temperature
zone instantaneously (in much less than the time spent in that zone). Consequently, the PCR time
in flow-through microPCR is limited by the polymerase extension rate. As in a pipe-flow
problem, the retention time of PCR reagent mixture in flow-through system is a function of flow
8

rate, channel cross-sectional area, and channel length. Starting with 10 copies, Kopp et al.
demonstrated the fastest flow-through microPCR in 90 sec, the theoretical limit of DNA
amplification time (Kopp et al. 1998).
Microfluidic chips for PCR and isothermal assays are commonly fabricated by a number
of inorganic and organic materials such as silicon, glass, polydimethylsiloxane (PDMS),
26

polymethylmethacrylate (PMMA), polycarbonate (PC), SU-8, and cyclic olefin polymer (COP)
(Fiorini and Chiu 2005; Zhang and Xing 2007b). Due to specific properties of these materials,
single material or a hybrid of two or more materials are commonly selected for chip fabrication.
For example, silicon provides high temperature ramping rates due to its high thermal
conductivity. Glass provides clear window for optical detection. Also glass is suitable for
capillary electrophoresis due to the electro-osmotic-flow property in glass chips is suitable for
capillary electrophoresis. Commonly reported micro/microfluidic chips are either fabricated
either with silicon or glass or a hybrid of both. Microchips made of silicon or glass is generally
not disposable due to high cost of fabrication. Polymeric microchips are inexpensive and
disposable. Hence they are becoming more common. Polymeric chips exhibit higher fabrication
flexibility, optical transparency, lower fabrication cost, and better biocompatibility. One
drawback of polymeric chips is the slow thermal response due to their lower thermal
conductivity (approximately 3-order of magnitude lower than silicon). Therefore, polymeric
chips may be more suitable for isothermal gene amplification, which does not require any
thermal cycling. Commonly applied microchip fabrication techniques are photolithography,
silicon bulk micromachining, soft lithography, injection molding, hot embossing, CO2 laser
ablation (Zhang and Xing 2007b). Bonding/sealing of microchip is important to contain PCR
reagent mixture and eliminate its evaporation during thermal cycling. A number of techniques
like anodic bonding (glass and silicon), hydrofluoric acid bonding (silicon and glass), adhesive
bonding (silicon, glass, and polymer), O2 plasma treatment (PDMS), lamination (polymer) have
successfully been applied (Zhang and Xing 2007b).
Here, a quantitatively evaluation of a number of parameters (e.g., thermal mass,
conductivity, channel cross-sectional area, and flow rate) that affect the nucleic acid
27

amplification time in static chamber and flow-through microPCR systems is performed. These
parameters can be optimized during the designing of microPCR chips to enhance system’s
performance and reduce PCR time. Some examples of microPCR systems with innovative
approaches and surface treatment of microchips are also briefly discussed. Examples of
integrated microPCR systems, which include sample processing and detection steps together
with the nucleic acid amplification step are discussed. Relatively, newer isothermal nucleic acid
amplification methods and the challenges associated with their translation on microchips are
discussed. As an example, a novel isothermal gene amplification method, LAMP, is discussed in
the context of simple sample processing and instrument-free diagnostic systems. Miniaturized
nucleic acid analysis systems are evaluated under the World Health Organization’s guidelines for
POC diagnostic devices, showing the potential of microLAMP systems towards simpler
integration of sample processing steps with other analytical steps in one device for POC
diagnostics. Here, the main aim is to critically evaluate the effect of various physical parameters
on the performance of microPCR and microLAMP systems for rapid and low-cost gene analysis
with the applications towards POC infectious disease diagnostics.
2.2 Static chamber microPCR systems
Static chamber microPCR chips with different materials, reaction throughputs, volumes,
starting copy numbers or amount of nucleic acids, heating and cooling rates, cycle numbers, and
amplification times has been reported (Table 2.1). The first micro-machined static chamber
microPCR system was reported by Northrup and coworkers in 1993 (Northrup et al. 1993a). The
system consisted of a silicon chip with a single 50 µL well, a computer controlled peltier unit for
thermal cycling, and silicone adhesive for well sealing. DNA amplification was completed in
approximately 1 h. A year later, a microPCR chip with 5-fold lower in volume (10 L) was
28

demonstrated (Wilding et al. 1994a). The chip was made of silicon with a single reaction
chamber covered by glass layer. Thermal cycling was performed by thermoelectric heater and
cooler. However, PCR time was increased to 1 h and 45 min. Higher PCR time with the 5-fold
reduction in the reaction volume may be due to lower heating rate (not reported). End products in
both the systems were detected off-chip by gel electrophoresis. Since then, static chamber
microPCR system has been replicated by numerous research groups, with the focus towards
reducing the amplification time, reaction volume, and cost, and integration of amplification,
detection, and analysis steps. Toward the first step, microchips were integrated with thin film
heaters to consolidate the amplification module. Three types of heating assemblies are
commonly applied to static chamber microPCR systems. These are: i) external heater i.e.
customized or commercial thermal cycler (Liu et al. 2001a), ii) integrated heater by using
metallic thin film (Lee et al. 2004), and iii) non-contact heater by using high energy radiation
(e.g., infrared rays) (Oda et al. 1998b). Rapid nucleic acid amplification can be achieved on the
microchip integrated with thin film heater or microchip with non-contact heater than on the
microchip placed on external heater. Additionally, heating source could be the bulkiest
component of the microPCR system considering the small dimension of the microchip and low
volume of the reaction solution. The reason for the differences in the PCR time for these heating
assemblies are quantitatively evaluated and described in the section below.
2.2.1 External heaters
External heaters are metallic blocks (e.g., copper or aluminum) attached to a peltier unit
for thermal cycling. Commonly these metallic blocks have cavities for placing PCR tubes. Static
chamber microPCR chip is directly placed on the thermal cycler with flat metallic plate (attached
with peltier) for temperature ramping. Due to slow thermal cycling rate of commercial thermal
29

cycler, it takes nearly same time to perform PCR in microchip as it will take in plastic vials. The
only advantage of performing PCR in microchip on commercial thermal cycler is the reduction
in the reaction volume applied to the microchip. Some of the reasons of using external
heaters/thermal cyclers with static chamber microPCR chip are; i) fabrication of thin film heater
on microPCR chip requires sophisticated facilities and instruments and is an expensive process,
ii) temperature control of thin film heaters is achieved by a combination of proportional-integraldifferential (commonly with LabVIEW software from National Instruments), which requires the
knowledge of electrical circuitry, and iii) fluorescence microscopes with temperature control
stage and customized software are available for real-time detection of PCR on chip. Therefore, a
number of researchers have applied commercial or custom-made thermal cyclers to perform PCR
in microchips.

30

Table 2.1 Selected parameters of static chamber microPCR systems
Chip Material

Reaction Volume
(nL)

Copy/amount

Heating source
(thermal cycler)

Polycarbonate
Glass

4

20,000
12,000

0.2 ng
0.12 ng

Commercial
Commercial

Heating/cooling Cycle PCR
o
time
rate ( C/s)
(min)
25
223
25
223

Glass-PDMS

8

10,000

-

Commercial

-

Silicon

1248

40

Commercial

PDMS-glass
PDMS

400
1

3
2000

0.4-8 DNA
copies
60 ng

Commercial
Custom-made

Glass-PDMS

1

2240

5 RNA copies

Silicon

1248

40

Glass-PDMS

3

Silicon

45

82.5

40

62

3/3
2/1.2

30

60
43

Custom-made

6/4

25

84

Custom-made

-

40

56

900

0.4 - 8 DNA
copies
0.017 ng

Custom-made

-

35

57

1

50,000

-

Custom-made

-

-

60

Glass-PDMS

4

3000

0.01 ng

Custom-made

-

35

87.5

Glass-PDMS
Polycarbonate
Silicon

100
1
1

120
10,000
10,000

Custom-made
20 DNA copies Custom-made
0.2 ng
Custom-made

4/6
-

40
45
35

65
101
105

Glass-PDMS

1

1750

11 ng

Custom-made

4.5

35

117

Glass-PDMS

9

5000

0.5 ng

Custom-made

-

40

122

31

Reference

(Liu et al. 2001b)
(Waters et al.
1998b)
(Zhang et al.
2010)
(Matsubara et al.
2004b)
(Liu et al. 2003b)
(Shin et al.
2003a)
(Kaigala et al.
2006)
(Matsubara et al.
2005b)
(Xiang et al.
2005a)
(Northrup et al.
1993b)
(House et al.
2010)
(Liu et al. 2009)
(Qiu et al. 2010)
(Wilding et al.
1994b)
(Prakash et al.
2006a)
(Ramalingam et
al. 2010)

Table 2.1
(cont’d)
Glass

1

0.08

15 DNA copies Custom-made

2/3-4

10

20

Glass-PDMS

72

0.45

34 RNA copies

-

-

30

45

PDMS

1

5000

0.1 ng

Custom-made

3.1/3.1

40

45

Glass-PDMS

1

100

10 cells

Custom-made

5/3

35

50

COC

1

50,000

1.25 ×106 cells

Custom-made

10/5

40

50

Polypropylene

1024

200

5 DNA copies

Custom-made

-

45

55

Glass
Glass-PDMS

4
1

3800
600

0.16 ng
24 RNA copies

Thin film
Thin film

2/1
10

30
35

35
44

Silicon-glass

10

100

100,000

Thin film

175/125

40

5.7

Silicon-PMMA
Glass

1
8

200
200

0.06 ng
5 ng

Thin film
Thin film

80/60
20/20

30
20

7
10

Silicon

1

3000

~60 ng

Thin film

11/11

30

13

Silicon
Silicon-glass

1
1

5000
1000

Thin film
Thin film

12.7/12.7
-

35
50

25
28

Glass

1

500

Thin film

11.5/5.6

50

19

Glass-PDMS

1

10,000

100,000 RNA
copies
12 cDNA
copies
-

Thin film

20.5/7

20

15

Glass-plastic

1

25,000

5 ng

Thin film

8/8

30

28

32

(Khandurina et
al. 2000a)
(Marcus et al.
2006a)
(Cady et al.
2005a)
(Beyor et al.
2009a)
(Sauer-Budge et
al. 2009)
(Dahl et al.
2006a)
(Pan et al. 2010)
(Kaigala et al.
2008)
(Neuzil et al.
2006a)
(Lee et al. 2004)
(Lagally et al.
2001)
(Rodriguez et al.
2003)
(Wu et al. 2011a)
(Neuzil et al.
2010)
(Pipper et al.
2007a)
(Hsieh et al.
2005b)
(Shen et al.
2005a)

Table 2.1
(cont’d)
Glass-PDMSPMMA
Glass-PDMS

1

10,000

1000 ng

Thin film

20/10

20

18

1

10,000

0.06 ng

Thin film

20/10

25

22

Glass-PDMS
Silicon-glass
Silicon-glass

4
1
4

390
240
3000

10 RNA copies
0.24 ng
1.2 ng

Thin film
Thin film
Thin film

15/10
20/20
6/3.5

30
35
30

27
22
30

SU-8

1

25,000

Thin film

20/11

40

30

Silicon-glass

1

1000

Thin film

20

40

20

Glass

5

100,000

28600 DNA
copies
278 DNA
copies
1000 ng

Hot air

-

30

15

Silica

1

62,500

0.25 ng

Hot air

-

25

8

Glass

1

28,000

5 ng

Infrared rays

10/20

17

14

Silica

1

160

593

Infrared rays

65/20

10

11.5

Glass-PDMS

1

550

15-45 ng

Infrared rays

-

30

11

Polyimide

1

1700

69,000 DNA
copies

Infrared rays

10/10

15

4

.

33

(Huang et al.
2006)
(Liao et al.
2005a)
(Liu et al. 2006a)
(Pal et al. 2005)
(Zou et al.
2005b)
(Wang et al.
2006)
(Cho et al.
2006b)
(Wittwer et al.
1989b)
(Swerdlow et al.
1997a)
(Oda et al.
1998a)
(Huhmer and
Landers 2000a)
(Easley et al.
2006b)
(Giordano et al.
2001b)

For example, Liu et al. reported a polycarbonate microPCR chip with a volume of 20 µL
integrated with capillary electrophoresis for integrated gene analysis (Liu et al. 2001a). DNA
amplification was performed in 223 min for 25 cycles by thermally cycling the microchip on a
commercial thermal cycler. Waters et al. designed a glass microchip with 4 reaction wells each
with a volume of a 12 µL (Waters et al. 1998a). Similarly, temperature cycling was performed on
a commercial thermal cycler providing amplification in 223 min for 25 cycles. Kaigala et al.
designed a 2 µL glass-PDMS microPCR chip for the detection of BK virus (Kaigala et al. 2006).
Thermal cycling was performed on a custom made thermoelectric system with a heating and
o

o

cooling rate of 5 C /s and 3 C /s respectively, leading to an amplification time of 84 min in 25
cycles. Later, the same group integrated microPCR chip with capillary electrophoresis to achieve
a $1000 portable system for gene amplification and detection (Kaigala et al. 2008). Increase in
o

heating rate to 10 C /s and a 4-fold reduction in the reaction volume, provided a faster
amplification in 44 min for 40 cycles. PDMS is an inexpensive and commonly applied
elastomeric polymer for microchip fabrication. However, the water diffusion and vapor loss
property of PDMS at elevated temperature is an issue for low-volume applications. To address
this problem, a protocol of implanting polyethylene barrier in PDMS was developed, which
eliminated the evaporation problem (Prakash et al. 2006b). Inter-diffusion of the dissolved
polyethylene into PDMS structure resulted into irreversible co-adhesion. Approximately, 3-fold
reduction in solution evaporation was achieved in the polyethylene implanted microchip.
Another solution of covering PDMS surface with parlene film drastically reduced the solution
evaporation (Shin et al. 2003b).
For high throughput DNA amplification, Matsubara et al. developed a nanoliter array
based microPCR chip (Matsubara et al. 2004a). A total of 1248 reactions wells each with 40 nL

34

2

volume were fabricated on a 1×3 inch silicon chip. A commercial thermal cycler was employed
for temperature cycling. In a following report, a miniaturized thermal cycler was developed for
PCR in the same 1248 reaction well microchip (Matsubara et al. 2005a). TaqMan assay was
applied for the amplification of 3 different serotypes and 5 genes of E. coli. Amplification was
achieved in 62 min for the system with commercial thermal cycler and 56 min for the system
with miniaturized thermal cycler. Heating/cooling rates of these thermal cyclers were not
reported. Despite the higher thermal conductivity of silicon than polymers, higher amplification
time might be due to the higher heating and cooling rates of the cycler. Dahl et al. developed a
1024 open well-based polypropylene chip with a volume of 200 nL per reaction for the
amplification of RNase P gene from human genomic DNA and high throughput gene expression
studies (Dahl et al. 2006b). Thermal cycling was performed on a custom-made cycling system
with a DNA amplification time of 55 min for 45 cycles. Liu et al. developed a PDMS-glass
microPCR chip for 400 independent reactions each with a volume of 3 nL per reaction (Liu et al.
2003a). Microchip was placed on a commercial thermal cycler with heating and cooling rates of
o

3 C /s. A 294-bp segment of the human -actin cDNA fragment was amplified in 60 min.
Marcus et al. demonstrated 72 real-time PCR reactions each with 450 pL volume in a glassPDMS microchip (Marcus et al. 2006b). Thermal cycling of the chip was performed on flat plate
thermal cycler leading to PCR time of 45 min. Analysis of the above examples show that only
reduction in PCR reagent mixture volume does not reduce PCR time. High PCR time in the
reported static chamber microPCR systems was due to the low heating and cooling rates of the
applied thermal cyclers and low conductivity of glass or polymeric microchips.

35

2.2.2 Integrated heaters
Integrated heaters are thin metallic layers ranging from fraction of micrometers to several
millimeters in thickness directly fabricated on the back-side of microPCR chip. In some
examples, thin films are fabricated on a substrate (e.g., glass or silicon), which is directly placed
in the contact of microPCR chip. Platinum, gold, aluminum, tungsten, indium tin oxide is
commonly used as materials for heaters. Thin film heaters are fabricated by physical vapor
deposition, sputtering, and laser deposition techniques. Integration of microchip with thin filter
heater provides efficient heat transfer to the PCR reagent mixture. Improper heating leads to
irreproducible thermal profiles with low PCR specificity and non-reproducible DNA
amplifications. Several researchers have applied finite element models (e.g., Coventor Ware,
ANSYS, FLUENT) to analyze the thermal response of thin film heaters (Lin et al. 2000; Zou et
al. 2005a; Dinca et al. 2009). These studies were helpful in optimizing the size, shape, position
of thin film heaters, resulting in uniform temperature distribution on the microchip. A number of
studies have demonstrated that thin film heaters provide higher heating/cooling rates than
commercial thermal cyclers, leading to rapid gene amplification (Zhang and Xing 2007a).
For example, Neuzil and coworkers demonstrated nucleic acid amplification only in 5.7
min, which is the fastest till date (Neuzil et al. 2006a). A silicon chip was fabricated in the form
of a cantilever with a disc at the end. A microscope glass cover slip was placed on the top of the
disc integrated with a thin film heater and a temperature sensor. A new concept of virtual
reaction chamber, where the mineral oil covers the PCR reagent mixture was applied to perform
PCR. A 100 nL PCR reagent mixture encapsulated with mineral oil was placed on a disposable
glass chip for thermal cycling. This type of arrangement provided an extraordinary heating and
o

o

cooling rates of 175 C /s and 125 C /s respectively. Rapid amplification was due to the low

36

thermal mass of the system, especially the thin film heaters as suggested by the authors. High
throughput was achieved by running 10 PCR reactions simultaneously. Real-time nucleic acid
amplification was monitored by placing the chip on a microscope attached to a photomultiplier
tube. Amplicons were analyzed by capillary electrophoresis to confirm the performance of PCR
at this high thermal cycling rate. It was shown that microPCR was highly specific. Lee et al.
fabricated a sub-microliter volume static chamber microPCR chip with very high thermal
response or low thermal mass (Lee et al. 2004). Microchip was composed of silicon membrane
base covered by PMMA top, etched with fluidic networks. Silicon membrane was fabricated by
the bulk-micromachining of silicon. Platinum heaters and sensors were placed on the silicon
o

o

membrane. Heating rate of 80 C /s and a cooling rate of 60 C /s was achieved. DNA
amplification in a 200 nL chamber was demonstrated in only 7 min with 30 cycles. Combination
of low thermal mass of heating membrane, efficient heat transfer in silicon, and low PCR
reaction volume resulted into faster heating and cooling and rapid DNA amplification. Lagally et
al. developed a glass-based microPCR chip with 8 parallel 200 nL PCR chambers integrated with
platinum thin film heater, and capillary electrophoresis channels (Lagally et al. 2001).
Temperature control was accomplished through a proportional/ integral/derivative (PID) module
0

with the LabVIEW program. Heating and cooling rates of 20 C /s were achieved. Multiplex
microPCR was carried out to amplify a specific region of X and Y chromosomes from human
genomic DNA in 10 min. After amplification, the sample was detected in the integrated capillary
electrophoresis channels. Similarly, Rodriguez et al. reported a silicon microPCR chip integrated
with glass-based capillary electrophoresis channels (Rodriguez et al. 2003). Heating and cooling
o

rates of 11 C /s were achieved. Moreover, 30 cycles of micro-PCR were completed in 13 min.

37

Hsieh et al. developed a battery-operated portable microPCR system integrated with glass-PDMS
chip and platinum thin film heater (Hsieh et al. 2005a). Temperature was measured by an
o

integrated sensor and controlled by a microcontroller. Heating and cooling rate of 20.5 C /s and
o

7 C /s was achieved. A high reaction volume of 10 µL was used. A 276 bp lytA gene for
Salmonella pneumoniae was amplified in 15 min. Detection of PCR product was carried out
separately by gel electrophoresis. Due to low power consumption (1.2 W/s) of controller, this
system could operate for 10 h with one time charging. Generally, microPCR chips are disposed
after DNA amplification to avoid any contamination by the amplification product. Disposing
microPCR chips attached with thin film heaters is not a very cost-effective option. Shen and
coworkers addressed this problem by designing a 25 µL microPCR reaction chamber
sandwiched between a glass slip and a plastic cover (Shen et al. 2005b). Glass chip and plastic
cover was separated by using a double sided adhesive tape between the glass and plastic cover.
Thin film heater was fabricated on the backside of glass slide. A reasonable heating and cooling
o

rate of 8 C /s was achieved. A 244 bp gene fragment of hepatitis C virus was amplified in 28
min. The above examples show that DNA amplification can be achieved faster in static chamber
microPCR chips integrated with thin film heaters as compared to commercial or customized
thermal cyclers. Applying thin film heaters to static chamber microPCR chips reduces PCR time
by enhancing the heat transfer and increasing the heating/cooling rates (Khandurina et al. 2000b;
Liao et al. 2005b; Huang et al. 2006; Liu et al. 2006b). Moreover, lower reagent volumes,
application of heat conductive chip materials, and precise temperature control are other critical
factors for rapid gene amplification.

38

2.2.3 Non-contact heaters
A disadvantage associated with contact heaters (e.g. peltier and thin film) is that a certain
amount of thermal mass is added to the system, which slows down the thermal cycling. Only an
effective contact between the heating source and PCR reaction vessel leads to efficient heat
transfer resulting into reproducible PCR with high yield. Heat transfer problems associated with
heating blocks can be eliminated by applying non-contact heating approach. Non-contact heating
options are applied to achieve faster heating/cooling and rapid DNA amplification. Some
examples of non-contact heating include heating PCR reagent mixture by hot air, infrared (IR)
light, laser, induction, microwave (Zhang et al. 2006). Hot-air and IR-based heating has been
commonly applied to microPCR systems than other non-contact heating options.
Hot air based thermal cycling of PCR reagent mixture was initiated by Wittwer and
coworkers (Wittwer et al. 1989a). PCR reagent mixture contained in glass capillary was heated/
cooled by rapidly switching the air streams at desired temperatures. A 536-bp beta-globin
fragment of human genomic DNA was amplified in 15 min. Swerdlow et al. demonstrated a hotair assisted PCR in the injection valve of high performance liquid chromatography (Swerdlow et
al. 1997b). All the necessary steps, including loading of PCR reaction mixture, amplification,
and detection by capillary electrophoresis was performed in 20 min with a PCR time of 8 min.
Hot air heats the PCR reagent mixture and the reaction vessel non-selectively, which is not an
efficient utilization of heat/power.
IR source is another effective non-contact heating option. Water absorbs IR light in the
wavelength range of 900-4000 nm, where the excitation of vibrational bands of water molecules
results in bulk heating of solution. Due to this property, IR source heats only the PCR reagent
mixture and does not heat the microchip and other parts depending on the selected material.

39

Moreover, IR light can be manipulated by applying lens and filters for focusing and eliminating
wavelengths (ultraviolet) that could be absorbed by other materials and could also interfere with
PCR. IR-based non-contact heating of microchamber was initiated by Landers and coworkers
(Oda et al. 1998b). Glass microchamber with a total volume of 28 µL was fabricated by
borosilicate glass. A tungsten lamp was applied as an IR radiation source for heating and
o

o

solenoid-gated compressed air source for cooling. This combination provided 10 C/s and 20 C/s
of heating and cooling rates respectively. The total PCR time including pre-cycle denaturation,
cycling time and post-cycle extension time was 14 min. In another report, IR-mediated thermal
cycling of 160 nL of PCR reagent mixture in fused-silica capillary was demonstrated (Huhmer
and Landers 2000b). Nano-liter volume of PCR reaction mixture required only 10 PCR cycles to
produce detectable amount of amplified products by capillary electrophoresis. High heating
o

o

(65 C /s) and cooling rates (20 C /s) were achieved. Based on IR-mediated heating, Giordano et
al. demonstrated the fastest PCR only in 240 s.(Giordano et al. 2001c) PCR reaction was
performed in a polyimide microchip with a reaction volume of 1.7 µL. Polyimide is a useful
material for IR heating due to its high optical transmission in IR range and high glass transition
o

temperature (350 C). For accurate temperature control, a thermocouple was placed inside the
reaction chamber. Taq polymerase inactivation was observed due to the direct contact of metal
with the PCR reagent mixture. Addition of 0.75% polyethylene glycol in the PCR reagent
mixture eliminated the problem of polymerase inactivation. These non-contact heating
assemblies do not contribute any thermal mass to the microPCR systems resulting into faster
thermal cycling and DNA amplification. IR-sources (e.g., tungsten lamp) applied in the above
examples were bulky and could not be a part of a portable/handheld microPCR system.

40

Miniaturized IR lasers are commercially available, which should satisfy the requirements of
efficiency and compactness and can be used in integrated microPCR systems.
2.2.4 Analysis of static chamber microPCR systems
The main objective of this analysis is to understand the key factors responsible for rapid
DNA amplification in static chamber microPCR systems. Static chamber microPCR systems
discussed above applied microchips fabricated with different materials, reaction volumes,
starting DNA copies or cells, external and non-contact heaters, integrated heaters with different
materials. As a result these microPCR systems have different heating/cooling rates and nucleic
acid amplification times. A general comparison of the above examples shows that PCR time is
related to heating/cooling rates. Additionally, heating/cooling rates are dependent on the volume
of PCR reaction mixture, conductivity of microchip material, and type of heater. For example,
Neuzil et al. applied the reaction volume of 100 nL, conductive silicon chip integrated with thin
film heater, and achieved the highest heating/cooling rates, leading to the fastest PCR in 5.7 min
(Neuzil et al. 2006b). On the other case, Liu et al. used a 20 µL PC chip on a commercial thermal
cycler and achieved the slowest PCR in 223 min (Liu et al. 2001a). Although the heating/cooling
o

rates of the commercial thermal cycler was not reported but it might be in the range of 1-2 C/s.
Therefore, the type of heaters applied to microchip has major effect on PCR time. Static chamber
microPCR chips heated with external heaters and thin film heaters from selected publications (xaxis) are plotted with respect to their reported PCR time (y-axis) (Figure 1.1). PCR time for
microchips heated with thin film heaters is in the range of 5 - 30 min, and 1 - 3 h for microchips
heated with external heaters (commercial or custom-made) respectively. This qualitative
comparison shows that microchips heated with thin film heaters have shorter PCR time than

41

microchips heated with external heaters. Also microchips with thin film heaters have higher
heating/cooling rates than the microchips with external heaters. Heating rates and PCR times are
plotted for the selected examples of static chamber micro-PCR chips heated with external heaters
and thin film heaters (Figure 2.1).

Figure 2.1 Comparison of PCR time for static chamber microPCR chips with thin films (black
bars) and external heaters (shaded bars)

Higher heating rates were achieved for microchips integrated with thin film heaters.
Lower heating rates were obtained for microchips placed on external heaters due to the slow
thermal cycling of commercial thermal cyclers. Also rapid amplification was achieved on the
microchips integrated with thin film heaters due to faster thermal cycling. Generally, external

42

o

heaters/thermal cyclers have heating rates of 1 - 2 C/s (Wittwer et al. 1989a), whereas thin film
o

o

heaters can be easily heated from 10 C/s to 175 C/s (Shen et al. 2005b; Neuzil et al. 2006b).
PCR time is inversely related to heating/cooling rates. However, it cannot be directly predicted
based on the heating/cooling rate. PCR time is directly dependent on the time needed for each
cycle. Cycling times can vary between seconds to few minutes per temperature cycle. During the
thermal cycling, time is required for bringing the temperature of the microchip to the cycle
equilibrium temperature and efficiently transferring heat to the PCR reagent mixture.
Temperature cycling rate is dependent on the thermal mass and conductivity of static chamber
microPCR chip. Therefore, rapid PCR can be achieved by a microPCR system using a
temperature conductive microchip with low thermal mass. In an study, it was observed that
replacing the plastic vials by glass capillaries and heating blocks by thin aluminum
compartments reduced the PCR time to 15 min (Wittwer et al. 1989a). It was proposed that rapid
amplification was achieved by reducing the thermal mass of the PCR system. Thermal mass,
which is a measure of the ability of the material to store thermal energy, can be described by the
multiplication factor of ρ. Cp. V, where ρ is the density of material, Cp is the specific heat, and V
is the volume of system. MicroPCR chips with low thermal mass would easily release the heat
during the cooling step, providing high thermal cycling rate. For heat conductive microchips
(e.g., silicon), heat transfer between the heater and the PCR reagent mixture would be rapid. This
would save the time which would otherwise be added in the thermal cycling time, if low heat
conductive microchips (e.g., polymer or glass) are applied.

43

Figure 2.2 Effect of heating rates () on PCR times () for static chamber microPCR systems.
Inset shows the relation between thermal mass of thin film heaters and PCR time

In earlier discussions, thermal mass was qualitatively correlated with the PCR time. Here,
the numerical values of the thermal mass of heaters are calculated and correlated with the PCR
time (Figure 2.2, Table 2.2). Inset of Figure 2.2, shows the relation between PCR time and the
calculated thermal mass of selected examples. In a commercial PCR machine, thermal cycling is
achieved by an aluminum/copper block attached to a peltier unit. However, in static chamber
micro-PCR systems, metallic films of platinum, gold, and copper are commonly applied as
heaters. For comparison, dimension of an aluminum block was assumed to be 5 cm × 5 cm × 5
3

cm. The specific heat and density of aluminum is 385 J /g. K and 8.96 g /cm respectively.

44

Therefore, the calculated thermal mass of a 1120 g aluminum block would be 431.2 kJ /K.
Thermal mass of thin film heaters were calculated (Table 2.2) by their given dimensions.
Thermal mass of copper block was much higher than the calculated thermal mass of thin film
heaters. Due to their high thermal conductivity, thin film heaters could be heated very fast to the
temperature needed for PCR. Also due to their low thermal mass, they would release the heat
easily during the cooling step. Therefore very high heating/cooling rates could be achieved by
thin film heaters. However, PCR time would not only depend on the heating/cooling rate of
heater. It would depend on how efficiently the heat is transferred from the heater to the PCR
reagent mixture through the microPCR chip and gets released during the cooling step. Therefore,
it is important that the comparison of the thermal mass of the whole microPCR system
(microchip, PCR reagent, and heater) should be done with the PCR time.

Table 2.2 Calculated thermal mass of thin film heaters for static chamber microPCR systems
Reference

PCR
time
(min)
7

Heater
material

Heater dimension

platinum

1500 × 500 × 0.3

(Hsieh et al.
2005b)
(Liu et al. 2003b)

15

platinum

π × (1500) × 0.1

27

platinum

(Shen et al. 2005a)

28

copper

(5×10 ) × (1×10 ) ×
0.2
120000 × 55 × 35

(Lee et al. 2004)

3

(m )
2

4

4

Thermal mass (J
/K)
6.27 × 10
1.97 × 10
2.79 × 10
7.95 × 10

-7
-6
-4
-4

Physical dimensions of microchips and properties of the applied materials (e.g., adhesive
tape) of were reported for few examples of Table 2.2. For example, Lee et al. (Lee et al. 2004)
and Shen et al. (Shen et al. 2005b) have reported all the parameters about microchip, PCR
reagent, and heater. Lee et al. applied a 200 nL silicon chip (covered with PMMA film) with a
dimension of 1.5 mm × 0.5 mm × 0.25 mm (Lee et al. 2004). Shen et al. applied a 25 µL glass

45

chip with a dimension of 22 mm × 22 mm × 0.15 mm.(Shen et al. 2005b) Glass chip was
covered by an adhesive tape and plastic cover. However, the physical properties of adhesive tape
and plastic cover were not given. Therefore, for both the examples we have calculated the
thermal mass of PCR chips and reagents for comparison. Specific heat and density of silicon is
3

0.70 J/g. K and 2.3 g /cm respectively. Specific heat and density of borosilicate glass is 0.88 J/g.
3

K and 2.2 g /cm respectively. For Lee et al. (Lee et al. 2004), thermal mass of PCR reagent
mixture and microchip was 0.84 × 10

-3

J/K and 0.30 × 10

-3

J/K respectively. For Shen et al.

(Shen et al. 2005b), thermal mass of PCR reagent mixture and microchip was 0.11 J/K and 0.14
J/K respectively. The total thermal mass of microchip with reagent mixture for Lee et al. (Lee et
al. 2004) and Shen et al. (Shen et al. 2005b) was 0.001 J/K and 0.25 J/K respectively. Thermal
mass of the thin films were not included in the calculations due to their very low values. It shows
that the combined thermal mass of the microchip and PCR reagent volume was approximately
three orders higher than the thermal mass of heater. Therefore, for the microchip and thin film
heater combination, PCR time should be affected more by the thermal mass of microPCR chip
and reagent mixture than the thermal mass of thin film heater. However, for the microchip and
external heater combination, PCR time should be more effected by the thermal mass of heating
blocks rather than microchip or PCR reagent.
2.3 Flow-through microPCR systems
Flow-through microPCR chips are commonly fabricated in silicon, glass, or a hybrid of
both (Zhang and Xing 2007b). Few studies have also reported polymeric (PDMS, PMMA) flowthrough microPCR chips (Zhang and Xing 2007b). Flow-through microPCR chips have
serpentine channels, which pass through three constant temperature zones or heaters necessary
46

for PCR. Instead of being stationary in the chamber, PCR reagent mixture is pumped
unidirectionally from the channel inlet passing repetitively through these thermal regions and the
amplified product is collected at the outlet. Generally, syringe or peristaltic pumps are used to
precisely control the flow of PCR reagent mixture in the microchannels. As described earlier,
heat inertia of a flow-through microPCR system is only the thermal mass of PCR reagent
mixture. The temperature transition time depends on the solution flow-rate and the time to
achieve thermal equilibrium. Consequently, the PCR time is limited by the extension rate of
DNA polymerase and could be much faster than static chamber micro-PCR. For rapid DNA
amplification, flow-through microPCR systems requires the optimization of channel crosssectional area, solution flow rate, and channel length, which are compared in Table 2.3.
The concept of flow-through PCR was introduced by Nakano et al., based on gene
amplification in a loop of Teflon capillary passing through three oil baths maintained at three
different temperatures (Nakano et al. 1994) The length of the capillary in each temperature bath
and flow rate of reagent mixture was fixed to achieve the optimized time for denaturation,
annealing, and extension. DNA amplification was achieved in 12 min. Similarly, Curcio and
Roeraade developed a PCR system by coiling a 15 m long Teflon tube on copper heating blocks
maintained at three different temperatures (Curcio and Roeraade 2003) A new concept of high
throughput was demonstrated by introducing PCR reagent mixture plugs with water immiscible
organic plugs. A compact cylindrical thermal-cycling device for flow-through PCR was reported
by Park et al. with gene amplification in a 3.5 m long fused silica capillary coil passing through
three heating blocks (Park et al. 2003a). DNA amplification was achieved in 5.5 min with a
reagent mixture flow-rate of 5 µL /min. Although the footprints of these setups were not in

47

microscale but the concepts were helpful for the development of flow-through microPCR
systems.
Kopp et al. were the first to demonstrate the fastest flow-through microPCR system in 90
s (Kopp et al. 1998). Serpentine channels (40 µm × 90 µm × 2.2 m) for 20 cycles were fabricated
on biocompatible glass substrate. The glass chip was placed on three copper blocks maintained at
3 different temperatures. A syringe pump was used to inject PCR reagent mixture (10 µL) with
varying flow rates of 5.8 nL/s to 72.9 nL /s corresponding to PCR time of 18.8 min to 90 s
8

respectively. However, the DNA concentration (1×10 copies) in the PCR reagent mixture was
very high. Detection sensitivity of microPCR system is as important as speed because very low
dose (e.g. 10 to 100 cells for E. coli O157:H7) can be a serious health risk.(Nataro and Kaper
1998) The initial demonstration was rapid but not very sensitive and required relatively high
reagent volume. Later, Obeid et al. demonstrated an improved version (in terms of sensitivity
and reagent volume) of flow-through microPCR chip with lower volume.(Obeid et al. 2003a)
Also, the chip had the flexibility of selecting the cycle numbers. Therefore, a range of starting
6

8

DNA copies (2.5 × 10 - 1.6 × 10 ) was successfully amplified. Incorporating air plugs between
the PCR reagent mixtures increased the throughput (3 parallel reactions) of the system. In
addition, fabrication of an extra cycle for provided the conversion of RNA to complementary
DNA before the PCR cycle leading to the amplification of DNA and RNA. Improved sensitivity
6

(2.5 × 10 copies), lower reaction volume (2-7.6 µL), and 5 min of PCR time was achieved. In
order to avoid bulky copper heaters, Hashimoto et al. evaluated the thermal effects by varying
the PCR reagent mixture flow rates on a 20 cycle spiral channel polycarbonate chip (Hashimoto
et al. 2004a). Effect of flow rates (1-20 mm/s) on the temperature distribution along the

48

microchannel was simulated. Effect on the temperature transition and resident times of PCR
reagent mixture with different flow-rates were estimated for the optimization of experimental
7

condition. Amplification of 2×10 copies of a 500 bp DNA template was completed in 1.7 min.
Additionally, detection of single base pair mutations on a polycarbonate flow-through microchip
was reported by the same group (Hashimoto et al. 2006b). Three different assays (flow-through
PCR, ligation detection reaction, and hybridization) were performed on single microchip. DNA
amplification on this 30-cycle serpentine channel microchip was achieved in 18.7 min. This
example shows the possibility of integrating the pre/post-PCR genetic analysis steps on flowthrough microPCR system. Schneegaß and coworkers used a flow-through chip made of silicon
to improve the temperature transition (Schneegaß et al. 2001). To achieve a smaller footprint,
platinum heaters were fabricated on the backside of the chip. Regions made of silicon ensured
the efficient heat transfer from the heaters to the PCR reagent mixture. However, silicon creates
the problem of temperature overlap between the channels due to its high thermal conductivity.
Incorporating air pockets between the channels reduced the heat transfer effect between
channels. Generally high surface area of channels leads to the adsorption of bio-molecules,
which affects the PCR efficiency (Shoffner et al. 1996a). The static and dynamic surface
treatment before and during the PCR step was applied for the optimization of PCR conditions.
Also the effect of flow rate on the amplification efficiency was studied. An optimum flow rate of
33 nL/s was demonstrated for a 25 cycle, 17.5 min flow-through microPCR. Xiaoyu et al.
reported a spiral channel microPCR chip integrated with platinum thin film heater (Xiaoyu et al.
2005).

The chip was a hybrid of PDMS and glass with a reagent volume of 42 µL.

Amplification of a 450 bp segment of E. coli HB101 was achieved in 30 min.

49

Table 2.3 Parameters for systems with flow-through design including flow-through microPCR systems
Factors/Reference

(Kopp et al. (Obeid et al.
1998)
2003b)

(Park et al.
2003b)

(Hashimoto et al. (Schneegaß et al.
2004b)
2001)

(Hashimoto
et al. 2006a)

(Sadler
et al.
2003b)

(Chou et al.
2002b)

Time for PCR (min)
Number of cycle
Flow rate (nL/s)
Cross-sectional area
2
(µm )
Channel length (m)
Volume of PCR
solution (µL)
Target copy

1.5
20
72.9
3600

5
20
21
5181

5.5
33
83.33
7850

1.7
20
22.5
7500

17.5
25
33
19625

18.7
30
6.67
5000

27
40
325
250,000

40
30
250
250,000

2.2
10

3.43
7.6

3.5
50

---------

1.512
33

1.57
0.168

.....
24

----19

-----

2 × 10 -1 × 10

------

------

------

-----

Polycarbonate

Glass

LTCC

LTCC

spiral loops

serpentine

Polycarbona
te
serpentine

serpenti
ne
Ag-Pd
thin
films
-----

serpentine

8

1 × 10

6

2.5 × 10 -1.6 ×

7

8

8

Substrate

Glass

10
Borosilicate glass

Design

serpentine

serpentine

Fused silica
capillary coils
helical

Thermal source

Copper
blocks

Copper blocks

Copper blocks

Resistive heaters

Platinum thin film
heaters

Surface treatment,
type

Dichlorodi
methylsilan
e, dynamic

Dichlorodimethyl
silane, static

Trimethyl
chlorosilane/D
MF/imidazole,
static

Bovine serum
albumin (BSA),
static

Hexamethyldisilan
e/ BSA, static and
dynamic

Film
resistance
heaters
No
treatment

Screen
printed AgPd films
-----

LTCC: Low temperature co-fired ceramics, PI: Proportional Integral PID: Proportional Integral Derivative, CFD: Computational Fluid Dynamics,
Ag-Pd: Silver-Palladium

50

Chou et al. demonstrated the application of a new substrate material, low temperature cofired ceramics for flow-through microPCR chip (Chou et al. 2002a). Screen printed silverpalladium paste was used as heaters. Computational fluid dynamics simulation was done to
optimize heater input power, solution flow rate, sensor placement, and air gap size for thermal
isolation. With a solution flow rate of 250 nL/s, a 20 cycle chip amplified a 209 bp DNA
template in 40 min. PCR products of flow-through microPCR chips are commonly analyzed off
the chip by gel electrophoresis. Development of optically transparent flow-through microPCR
chips could be useful for real-time detection of PCR products. Although, flow-through PCR
chips made of glass is transparent but the overall system is not optically clear due to the
associated heaters. Thin film of indium tin oxide is conductive and transparent and can be easily
coated on any substrate by electron beam evaporation, physical vapor deposition, and sputtering
technique. Sun et al. had used quartz glass chip coated with thin film of indium tin oxide heaters,
which made their system completely transparent (Sun et al. 2002). For this study, temperature in
the microchannels was initially calibrated by measuring the fluorescent spectra of temperature
indicator dye under the flow conditions. It was confirmed that indium tin oxide heaters provided
a local homogeneous temperature distribution on the chip. A 450 bp E.coli gene was amplified in
22.5 min. Recently a group have demonstrated a real-time flow-through micro-reverse
transcriptase PCR system for the detection of viral phage (Pjescic et al. 2010b). Reverse
transcriptase microPCR was performed in serpentine channels fabricated by xurography and
sandwiched between two glass substrates. LC green, a DNA intercalating dye was applied for the
8

florescence imaging of chip by a CCD camera. Approximately 1 ×10 copies of viral phage were
amplified in 15 min with 23 cycles. A single fluorescence image of the chip during the
amplification was sufficient to extract real-time amplification curves. Increasing the exposure

51

time of CCD camera from 125 ms to 1000 ms increased the signal-to-noise ratio (SNR) by 10fold.
Fabrication of microfluidic channels in glass or silicon requires sophisticated tools and
overall is an expensive process. A cost-effective version of flow-through microPCR chip was
developed by Kim at al. (Kim et al. 2006). Microfluidic channels were fabricated on a PDMS
chip by soft lithographic techniques. PDMS chip was laminated with glass and heated by
aluminum blocks. PDMS is hydrophobic in nature and generally adsorbs the enzymes used in
PCR reagent mixture. The effect of extension time and the treatment of microchannels by
polyvinylpyrrolidone on the PCR efficiency were studied. MicroPCR chip required only 8 min to
amplify a 430 bp DNA template. Another low cost flow-through microchip made of PMMA
substrate was reported by Sun et al. (Sun et al. 2008). PMMA is a transparent plastic with high
o

mechanical stability and moderate thermal transition temperature (105 C). It is suitable for costeffective fabrication of microfluidic channels by CO2 laser ablation or hot embossing. Sealing of
polymeric chips is a common problem (Grodzinski et al. 2001). For comparison of heat tolerance
o

o

at 95 C, PMMA chip was bonded with a PC cover (glass transition temperature 150 C) and
PMMA cover. It was found that PMMA-PC sealing had better performance than PMMA-PMMA
o

sealing. Due to lower glass transition of PMMA than PC, at 95 C, PMMA-PMMA bonding was
deteriorated leading to air bubbles and leakages. A number of solution flow rates (125 nL/s to
500 nL/s) for PMMA-PC flow through microchip were tested. For the amplification of a 500 bp
DNA template, 26 cycles were completed between 8.5 - 34 min.
The above examples show that rapid gene amplification between 90 s to 10 min is
possible in flow-through microPCR systems. Due to the applications of polymeric microchips,

52

flow-through microPCR systems can also be cost-effective. One major drawback of flow6

8

through microPCR systems is the low detection sensitivity in the range of 10 to 10 copies. On
2

the other hand, static chamber microPCR could be sensitive to approximately 10 copies (Cho et
al. 2006a). Also, surface modification of microchip is needed due to adsorption of PCR reagent
mixture in the channels of high surface area to volume ratio. Additionally, integration of parallel
and real-time amplification steps is difficult to incorporate in these systems. It might due to all
these reasons, including the bulky pumps to drive PCR reagent mixture, no commercial flowthrough microPCR system is yet available, even though its rapid nature was proven a decade
ago.
2.3.1 Analysis of flow-through microPCR systems
In the flow-through microPCR chip, nucleic acid amplification is achieved by driving the
PCR reagent mixture through three thermally isolated reaction zones. As the PCR reagent
mixture attains the temperature equilibrium with the heating zones instantaneously and thermal
cycling is not required, PCR time is greatly reduced (Hashimoto et al. 2004a). However, there
are other parameters e.g., PCR reagent mixture flow-rate, channel cross-sectional area, and
channel length that directly influences the PCR time in flow-through microPCR system. The
relation between PCR time and the ratio of cross-sectional area of the channel to the flow-rate of
PCR reagents is plotted in Figure 2.3. PCR times and the flow-rates were used as reported in the
publications. Cross-sectional areas of the channels were calculated by the reported channel
dimensions.
Comparison of the selected flow-through microPCR systems (Table 2.3 and Figure 2.3)
indicates that the PCR time is linearly dependent on the inverse of the flux of the PCR reagent

53

mixture. Therefore, rapid amplification in a flow-through microPCR system can be achieved
either by increasing the flow-rate and/or decreasing the channel cross-sectional area. For
example, the system presented by Kopp et al.(Kopp et al. 1998) had the flow-rate of 72.9 nL /s
2

with the minimum cross-sectional area of 3600 µm . Due to the minimum value of the ratio of
2

cross-sectional area and flow-rate (49 µm -s/nL), PCR was achieved in 90 s. Other factors for
rapid PCR were relatively smaller volume of PCR reagent mixture and lower cycle number.
Chou et al. (Chou et al. 2002a) and Sadler et al. (Sadler et al. 2003a) applied flow rates of 325 nL
/s and 250 nL /s, which lead to PCR in 27 min and 40 min respectively. In their cases, the
applied flow-rates and channel cross-sectional areas were highest among all the examples.
Therefore, their ratio of cross-sectional area and flow-rate were higher than Kopp et al. (Kopp et
al. 1998), leading to high PCR time. Ratio of cross-sectional area and flow-rate for other
examples were in between to Kopp et al.(Kopp et al. 1998), Cho et al.(Chou et al. 2002b) and
Sandler et al. (Sadler et al. 2003b).

54

Figure 2.3 Relationship between PCR time and the ratio of cross-sectional area and flow-rate.
Symbols and solid line corresponds, respectively to the data from studies reported on flowthrough microPCR systems and best fit line
2.4 MicroPCR systems with innovative concepts
Miniaturized PCR systems based on simpler concepts such as convective, thermal
gradient, and oscillatory microPCR have the potential for low-cost disease diagnostics. The
concept of Rayleigh-Benard convective PCR was first reported by Krishnan and coworkers in
2002 (Krishnan et al. 2002). Rayleigh-Benard convective PCR applies the principle of natural
convection to amplify DNA in a very simple format. Rayleigh-Benard cell has a cylindrical
o

o

cavity filled with PCR reagent mixture and covered by two plates at 61 C and 97 C covering
the top and bottom. Due to the temperature difference across the cavity, a static temperature
gradient is generated, which repeatedly circulates PCR reagent mixture through temperature
55

zones necessary for amplification. Increasing the convective flow velocity in the cavity reduces
the PCR time due to the faster solution cycling. It was observed that flow velocity increases with
increasing temperature difference and cylinder dimensions. A 295 bp segment of the human actin gene was amplified by this system. Sample volume was 35 L and PCR time was ~90 min.
In a different publication the same authors performed flow visualization (tracking the motion of
microspheres) and computational studies for convective PCR (Krishnan et al. 2004a). Multi-well
arrays (15 L each) were fabricated from a Plexiglas block for multiplex convective PCR. A 191
bp fragment associated with membrane proteins of influenza-A virus and a 295 bp segment of
the human -actin gene were amplified in 15 min. Recently, Agrawal et al., presented a pocketsized convective PCR thermal cycler (Agrawal et al. 2007). A triangular loop of fluoropolymer
tubing was heated by thermoelectric heaters from two inclined sides. The benefit of using
o

fluoropolymer tubing is that it is stable at high temperature (205 C) and compatible to PCR. The
versatility of this device was demonstrated by amplifying a multiplex mixture of targets from
five different respiratory viruses in 50 min. The hardware (single heater, heat control circuit, 2
AA batteries) costs approximately $10 with few cents per reaction for the disposable tubing. This
is the demonstration of the cheapest prototype for nucleic acid amplification. This technology
has recently been adopted by Gahaga Biosciences and launched with a commercial name of
LavaAmp. Cost of this thermal cycler has not yet fixed but expected to be approximately $ 300
(personal communication). Some general advantages of convective PCR are simple hardware (no
moving parts) and possibility of multiplexing and integration. Currently the amplified products
are measured separately by gel or capillary-electrophoresis.
Continuous-flow thermal gradient PCR was initially reported by Crews et al. (Crews et
al. 2008) Thermal cycling was achieved by cycling PCR reagent mixture within a spatial

56

temperature gradient. Temperature gradient was created by heating the glass microchip (with
embedded serpentine channels) along one edge and cooling along the opposite edge. By using an
IR camera, a 3D temperature distribution was recorded and the temperature was optimized for
the experimental conditions. By this innovative concept of continuous-flow thermal gradient, a
40 cycle PCR was achieved in less than 9 min. Recently, an example of thermal gradient-based
real-time flow-through microPCR system has been reported (Pjescic et al. 2010a). A 23 cycle
serpentine-shaped adhesive tape was sandwiched between two glass slides. The sandwiched chip
o

o

was placed between two aluminum heaters at 60 C and 90 C to generate a thermal gradient
profile for PCR. PCR reagent mixed with DNA intercalating due was pumped with a flow rate of
2 µL/min in the channel and imaged by a CCD camera under the illumination of light emitting
8

diode. A 110 bp of viral phase DNA template with 10 copies /µL was amplified in 15 min. The
major advantage of the single heating/cooling region is the small footprint of the microchip, no
thermal crosstalk between channels, simple hardware, and short PCR time.
Oscillatory-flow PCR was proposed and experimentally validated by Wang et al. (Wang
et al. 2005). Oscillating-flow PCR chip was fabricated by the silicon microfabrication technique.
The chip had straight channel with three independent temperature zones necessary for PCR. A
droplet of 1 µL PCR reagent mixture was pumped back and forth 35 times in an oscillatory-flow
mode through the straight channel. Amplification of human papillomavirus DNA amplification
was achieved in 15 min. Numerical simulation had shown that less than 1 s is required for the
droplet to maintain thermal equilibrium with the set temperature. In this study, PCR reagent
mixture was maintained at each temperature for 4 s with an additional 4 min of extension.
Münchow et al. demonstrated that a 40 cycle oscillatory-flow PCR could amplify a 372 bp
template in 5 min (Munchow et al. 2005). Functionally, an oscillatory-flow PCR is like a hybrid

57

of static chamber and flow-through PCR. It combines the flexibility of incorporating multiple
channels for high throughput applications, smaller chip size, and fast temperature transitions for
rapid PCR.
2.5 Surface treatment of microPCR systems
Surface chemistry plays an important role for the success of PCR in microPCR systems.
Due to small dimensions of microPCR systems, surface-to-volume ratio gets extremely high. For
example, the surface-to-volume ratio of a microsystem (characteristic length of 1 µm) is 10

6

times higher than a system with a characteristic length of 1 m. PCR reaction solution generally
consists of DNA polymerase, target DNA, primers, and salts (Saiki et al. 1988).

For an

optimized reaction, the concentration of PCR components should be maintained during the
course of the reaction. Miniaturization leads to non-specific adsorption of biomolecules on the
surfaces which change the concentration of the PCR components in the reaction mixture. This
phenomenon reduces the efficiency and efficacy of microPCR systems. MicroPCR chips are
mostly manufactured by silicon. The effect of native silicon induced PCP inhibition is widely
reported in literature. Coating of surfaces by PCR compatible materials or passivating agents
such as silicon oxide, BSA, polymers (poly vinyl pyridine and polyethylene glycol), silanes and
surfactants have shown considerable improvements in microPCR efficiency (Table 2.4). Based
on the application mechanism, surface treatment can be classified into two types; static and
dynamic passivation.
In static passivation, surface is coated with a PCR compatible substance before
performing PCR on the chip. Passivating agents strongly adhere to the surface by covalent
bonding. It is performed by filling and incubating the reaction chamber/channel (s) by a

58

passivation agent for a certain time followed by washing and curing the chip. Although it follows
a complex protocol but provide better performance. However, it is a time consuming (~24 h)
process. Covalent linked static coatings for silicon substrate are oxide deposition, polymer film
deposition, and silanization. Dynamic passivation occurs during the operation of microPCR chip.
Passivation agent is mixed with the PCR reaction mixture. During the PCR reaction, the
passivation agent preferentially binds on the inner surface of the microchip and prevents the
adsorption of PCR reaction components. The most frequently applied passivating agents for
dynamic coating are BSA, polyvinyl pyridine, polyethylene glycol, and non-ionic surfactant
(Tween).
Shoffner et al. studied the effect of native silicon on PCR inhibition (Shoffner et al.
1996b). Deposition of a thin layer of oxide or protein on silicon microchip slightly improved
PCR product yield. Felbel et al. suggested that a combination of silanization and BSA treatment
of silicon surface would be more effective than only one type of treatment (Felbel et al. 2004b).
Matsubara et al. also found that native silicon inhibits PCR (Matsubara et al. 2004c; Matsubara
et al. 2005c). Moreover, surface adsorption was predominant due to the high surface area to
volume ratio of nanometer-sized wells. Depositing an oxide layer followed by BSA treatment
improved the PCR yield as observed by the increment in the fluorescent signals. Krishnan et al.
studied the problem of non-uniformity of PCR reactions within the oxide coated silicon
microchannels (Krishnan et al. 2004b). The problem was due to the adsorption of positively
charged magnesium ions on the negatively charged silicon-oxide surface. Coating with Teflon
and proving higher concentration of magnesium ions in PCR solution showed improvements in
PCR efficiency.

59

Table 2.4 Examples of commonly used surface treatment strategies and their effect on microPCR
efficiency
Substrate
Silicon

Type of treatment
No treatment

Effect
PCR inhibition

Silicon

Silicon nitride

PCR inhibition

Silicon

Silicon oxide

Silicon

Silanization + BSA

Silicon
Silicon

Silanization + polyadenylic
acid (protein)
Silicon oxide + BSA

Glass

BSA

Glass

Polymers

Glass
PDMS

Silanization + Tween 20 +
polyvinyl pyridine
Parylene

SU-8

Silanization

SU-8

BSA

Polycarbonate

BSA

Pico TiterPlate
(Fiber)
Polyimide

BSA + Tween

Slightly improved
PCR product yield,
PCR inhibition
Increased specificity
and yield of PCR
product
Increased yield of
PCR product
Increased
fluorescence signal
and reduced PCR
inhibition
Reduced adsorption of
polymerase and DNA
Increased PCR
efficiency
Reduced adsorption of
enzyme and DNA
Eliminated bubble
formation, sample
evaporation, and
polymerase adsorption
Decreased the amount
of polymerase needed
and increased PCR
compatibility
Eliminated PCR
inhibition
Reduced adsorption of
polymerase and DNA,
enhanced PCR
compatibility
Eliminated PCR
inhibition
Reduced Taq
polymerase
inactivation

Polyethylene glycol

60

References
(Shoffner et al.
1996c)
(Shoffner et al.
1996c)
(Shoffner et al.
1996c)
(Felbel et al. 2004a)

(Shoffner et al.
1996c)
(Matsubara et al.
2005b), (Cho et al.
2006b)
(Khandurina et al.
2000a)
(Giordano et al.
2001a)
(Kopp et al. 1998)
(Shin et al. 2003a)

(El-Ali et al. 2004a)

(Wang et al. 2006)
(Liu et al. 2001b),
(Hashimoto et al.
2004b)
(Leamon et al. 2003a)
(Giordano et al.
2001b)

Giordano et al. compared the silanization process with different types of coatings such as
polyethylene glycol, polyvinylpyrrolidone, hydroxyethylcellulose, and polydimethylacrylamide
on the PCR yield of a glass microPCR chip (Giordano et al. 2001a).Untreated glass inhibited the
PCR process. Surface treatment with polymeric agents showed comparable results as silanization
together with being simple and less time consuming. Shin et al. applied a coating of inert
material called parylene on PDMS microchip (Shin et al. 2003c). This coating solved several
critical problems of bubble formation, sample evaporation, and protein adsorption for porous
PDMS substrate. El-Ali et al. observed that silanization of SU-8 chip (a negative photo resist)
increased the PCR product yield (El-Ali et al. 2004b). Liu et al. observed a reduced adsorption of
enzyme and DNA by coating the polycarbonate chip with BSA (Liu et al. 2001b). Leamon et al.
demonstrated that inclusion of 0.5% BSA and 0.05% Tween 80 in the PCR reaction solution
improved PCR efficiency (Leamon et al. 2003b). Channels of flow-through microPCR chips are
commonly treated with silanes, BSA, and nonionic surfactants (Table 2.3). In addition to surface
passivation, increasing the concentration of PCR components in reaction solution has also been
recommended.
2.6 Integrated microPCR systems
The aim of integrated microPCR system is to analyze real biological samples with a
sample-in-answer-out capability. So far the efforts to develop simple and rapid nucleic acidbased devices have focused predominantly on amplification and detection steps, with sample
processing remaining the bottleneck in the development of completely integrated systems.
Applications of miniaturized devices in clinical or environmental analysis require processing of
complex biological samples as whole blood, saliva, nasal aspirate, and environmental fluid on
the same device. The presence of inhibitory substances in biological samples like humic acids,
61

cell debris, particulates, metals etc. hamper microfluidic operations and biochemical reactions.
Generally, low initial target copies are present in biological samples, which requires target
enrichment step before amplification. Therefore, nucleic acid samples for microPCR are
commonly prepared off chip by applying established bench-scale procedures. Common steps for
sample processing includes, cell capture, cell concentration, cell lysis, nucleic acid extraction and
purification. The ability to integrate laboratory operations and automatic handling of biological
samples into a microPCR system is a key to the success of micro-total-analysis systems. The
examples discussed below can be divided into two major types, i) microPCR systems integrated
with amplification and detection modules, and ii) microPCR systems integrated with sample
processing, amplification, and detection modules.
Researchers at Lawrence Livermore National Lab developed a portable and a hand-held
device for the nucleic acid analysis. Both the devices were based on the concept of static
chamber microPCR. Example of the portable device is advanced nucleic acid analyzer
(Belgrader et al. 1998). The platform was made of silicon (container for plastic reaction tubes)
substrate heated by resistive heaters. It took 16 min to complete 10 PCR reactions
simultaneously. Real-time PCR analysis was performed on a separate laptop, as this device was
not capable to automatic data processing. Further developments lead to a hand-held device
named as handheld advanced nucleic acid analyzer.(Higgins et al. 2003) It was based on silicon
chamber and platinum heaters for the rapid heating and cooling of 4 plastic reaction tubes.
Amplification was achieved in 30 min. Moreover, results were analyzed within the device and
displayed on its screen. Currently, handheld advanced nucleic acid analyzer is marketed as BioSeeq through Smiths Detection for $ 31,000. The heating and cooling rates and reaction volume
o

for these devices were 1.5 C /s and 25 µL respectively, very similar to conventional PCR

62

machines. As the single PCR cycle was 23 s long, compared to approximately 240 s for
conventional PCR, rapid amplification was achieved. These systems can be further improved by
applying efficient thermal cycler, replacing plastic vials by highly conductive reaction
containers, and lowering the reaction volume. Pal et al. developed an integrated static chamber
microPCR system for genetic analysis of influenza virus (Pal et al. 2005). Microfluidic chip with
addressable valves and thermal components (heaters and sensors) were integrated on a one
system. RNA samples from the A/LA/1/87 influenza virus were amplified by reversetranscriptase PCR and detected by electrophoretic separation. MicroPCR was completed in 22
min. These examples used pre-processed nucleic acids for DNA/RNA amplification.
Integration of sample processing, amplification, and detection steps were also reported on
one system. For example Cady et al. integrated solid phase DNA extraction with real-time PCR
on a microfluidic silicon chip (Cady et al. 2005b). Silica-coated micro-pillars were used to
extract DNA from Listeria monocytogens cells. Temperature of the microchip was cycled by a
o

thermoelectric heater/cooler achieving a heating and cooling rates of 3.1 C /s. Approximately
4

10 cells of L. monocytogens were detected in 45 min. Legendre et al. reported a hybrid silica
bead/sol-gel based solid phase extraction and amplification of DNA from 600 nL whole blood
(Legendre et al. 2006). DNA amplification was achieved in 15 min by applying IR-based thermal
cycling. PCR products were detected by capillary electrophoresis. Easley et al. integrated a static
chamber microPCR with solid phase DNA extraction and detection by capillary electrophoresis
on the same device (Easley et al. 2006a). They demonstrated the performance of the device by
detecting Bacillus anthracis in blood sample and Bordetella pertussis in nasal aspirate. Starting
with sub-microliter sample volume, complete analysis required less than 30 min. DNA
amplification was achieved in 11 min by employing IR-based thermal cycling. Pipper et al.

63

reported a fully integrated system, which included sample processing, amplification, and
detection of avian influenza virus (H5N1) (Pipper et al. 2007b). Viral RNA was isolated directly
from throat swab sample, purified, and pre-concentrated using a droplet containing
superparamagnetic particles. The purified RNA was then released in a droplet, containing reverse
transcriptase PCR mixture. The droplet moved over the different temperature zones for thermal
cycling. The droplet acted as solid phase nucleic-acid extractor, pump, valve, mixer, and realo

o

time thermal cycler. Heating and cooling rates were 11.5 C/s and 5.6 C/s respectively.
Complete genetic analysis required only 28 min. An integrated microPCR system for the
detection of bacteria including lysis, DNA purification, PCR, and fluorescence readout has been
published recently (Sauer-Budge et al. 2009). Thermal cycling was achieved by a ceramic heater
o

o

6

(10 C/s) and air cooling (5 C/s). Detection of 1.25 × 10 cells of B. subtilis was performed in
approximately 40 min. Chen et al. developed a self-contained polycarbonate cassette for the
detection of B. cereus and HIV virus in saliva samples (Chen et al. 2010). All the necessary
liquid buffers and reagents were pre-stored in small pouches on the chip. All the steps including,
ell lysis, nucleic acid isolation, PCR, amplicon labeling, and detection by a lateral flow assay
were performed on the cassette. Amplification time was 28 min with a heating and cooling rate
o

o

3

4

of 5.2 C/s and 6 C/s respectively. The detection limit of this device was 10 -10 CFU/mL.
Beyor et al. demonstrated a glass-PDMS chip for cell concentration, purification, PCR, and
detection by capillary electrophoresis (Beyor et al. 2009b). The chip was used to detect E. coli
K2 and E. coli O157:H7 with a detection limit of 0.2 CFU/µL. Commercial nucleic acid analysis
systems have also been developed by microfluidic companies in a variety of formats and
functionalities, including ACLARA BioSciences, Fluidigm Corporation, Affymetrix, Agilent

64

Technologies, Alderon Biosciences, Roche Molecular Diagnostics, and Motorola Inc (Lee and
Hsing 2006). These miniaturized diagnostic tools would provide faster and accurate diagnosis,
significantly reducing the health care costs.
2.7 Isothermal nucleic acid amplification methods on microchip platforms
Isothermal amplification methods use the enzymes for in vivo synthesis of DNA/RNA at
a constant reaction temperature (Gill and Ghaemi 2008). A number of isothermal nucleic acid
assays have been developed with different types and numbers of enzymes and primers,
incubation temperatures, amplification times, and detection methods (Table 2.5). Isothermal
nucleic acid amplification assays such as loop-mediated isothermal amplification (Fang et al.
2010; Ahmad et al. 2011), multiple displacement amplification (Marcy et al. 2007), helicase
dependent amplification (Ramalingam et al. 2009; Mahalanabis et al. 2010), recombinase
polymerase amplification (Lutz et al. 2010; Shen et al. 2011), and nucleic acid sequence-based
amplification (Gulliksen et al. 2004; Tsaloglou et al. 2011) have also been translated on
microchip-based platforms. Unlike PCR, polymerase extension rate is the limiting factor in the
isothermal nucleic acid amplification assays. Therefore, detection times of these assays in
conventional detection formats (e.g., commercial real-time PCR instrument) are considerably
high. However, optimization of primer design and reaction, and application of highly fluorescent
Syto dyes and sensitive detectors would reduce the detection time to less than 30 min (Ahmad et
al. 2011; Kimura et al. 2011). Microchip-based isothermal nucleic acid amplification assays do
not require any sample or temperature circulation, which translates into simpler and less
expensive instrumentation.

65

Table 2.5 Selected examples of isothermal nucleic acid amplification methods
Method

Genetic
target

Reaction
temperature
o
( C)

Reaction
duration
(min)

Major Reaction
components

Reference

Loop-mediated
isothermal
amplification

DNA/RNA

60-65

30-90

1 enzyme (Bst
polymerase),4-6
primers

(Notomi et
al. 2000b)

Transcriptionmediated
amplification

DNA/RNA

60

140

2 enzymes (T7
RNA polymerase,
reverse
transcriptase), 2
primers

Helicase
dependent
amplification

DNA

65

75-90

3 enzymes (UvrD
helicase, DNA
polymerase I,
DNA-binding
protein), 2 primers

Recombinase
polymerase
amplification

DNA/RNA

37

20

3 enzymes (Bsu
polymerase, uvsX
recombinase,
DNA-binding
protein), 2 primers

Nucleic acid
sequence-based
amplification

RNA

41

105

3 enzymes (reverse
transcriptase, T7
RNA polymerase,
RNase H), 2
primers

(Hofmann et
al. 2005)

(Vincent et
al. 2004)

(Piepenburg
et al. 2006)

(Compton
1991)

2.7.1 MicroLAMP systems
Due to the simplicity and robustness of LAMP assays, microLAMP systems have
potential for POC disease diagnosis. LAMP is a relatively new gene amplification technique
initially reported in 2000 (Notomi et al. 2000a). Appealing features of LAMP include high
specificity, high sensitivity, speed, and high product yield. LAMP is performed at a moderate

66

o

o

incubation temperature between 60 C to 65 C. Major reaction components are an enzyme with
strand displacement activity (Bst polymerase) and four primers. Addition of two more primers
called loop primers further enhances the sensitivity and speed of LAMP reactions (Nagamine et
al. 2002). Positivity of LAMP reaction in PCR vial is confirmed by observing the turbidity (Mori
et al. 2001) or fluorescence (Chen and Ge 2010; Jia et al. 2010), either in endpoint or real-time
formats by incubating the vial in a water bath, disposable pocket hand warmers (Hatano et al.
2010), or commercial thermal cycler (Mori and Notomi 2009). High amplicon yield of the
LAMP reaction (500 ng /mL) simplifies the fluorescence detection due to increased signal
intensity in the presence of highly fluorescent dyes (Seyrig et al. 2010). Amplification times of
tube-based LAMP assays typically vary between 30-90 min, depending on the starting DNA
template (Seyrig et al. 2010).
Endpoint and real-time LAMP assays employed with turbidity or fluorescence-based
detection schemes have also been translated on microchips made of silicon and polymer
(Tourlousse 2010; Niemz et al. 2011). Positive endpoint microLAMP is confirmed either
visually (Fang 2010; Wu et al. 2011b) or by fluorescence microscopy (Matsubara et al. 2002;
Lam et al. 2008). Fluorescence (Abe et al. 2011; Liu et al. 2011) or turbidity (Fang et al. 2010)
increment during real-time microLAMP is commonly monitored by photodiodes. MicroLAMP is
further quantified by measuring the threshold time (Tt) for a selected cutoff of signal, where the
signal is above the background (Fang et al. 2010; Abe et al. 2011). Our group has optimized this
technique for real-time fluorogenic LAMP (using Syto-82 dye) on traditional real-time PCR
equipment and hand-held field deployable device (Gene-Z), both employing photodiodes
(Stedtfeld 2009; Tourlousse 2010). Recently, we have demonstrated that applying a CCD with
exposure control effect to real-time fluorescence microLAMP assays drastically increases the

67

SNR and reduces the Tt value in comparison to commercial real-time PCR instrument (Ahmad et
al. 2011).
2.7.2 Instrument-free LAMP systems
At present, microLAMP systems do not completely satisfy WHO criteria of equipmentfree/electricity-free operation. However, recent research has shown the possibility of equipmentfree microLAMP in future. For example, researchers at PATH (Seattle, WA) demonstrated
electricity-free LAMP by using calcium oxide heat source thermally in combination with a phase
change material (LaBarre et al. 2010; LaBarre et al. 2011). The phase change material is
o

designed to melt at 65 C with the heat generated by the reaction of calcium oxide and water.
This phase change material absorbs this heat and releases latent heat at relatively contact
temperature. It has been shown that the combination of 20 g of calcium oxide and 6.8 mL of
o

water provides sufficient heat that melts the phase change material and maintains 65 ± 0.75 C
temperature for 45 min. Fluorescence and turbidity-based endpoint LAMP assays (in 0.2 mL
PCR vials) targeting Plasmodium falciparum was evaluated on this chemical heater and also
compared against a commercial PCR instrument. Detection limit of Plasmodium falciparum on
both the systems was approximately 50 copies. Among the few limitations of this approach are
the slow warm up time (time required for the temperature of heater and reaction tube to reach
equilibrium) of 15 min, high incubation time for endpoint LAMP assay, and non-usability of
phase change material after single reaction. The warm up time could be reduced by applying the
chemical heater to micro chip due to its lower thermal mass. Also applying highly fluorescent
dye such as Syto-82 should further reduce the detection time of instrument-free microLAMP
assays.

68

WHO recommends that the POC diagnostic devices developed for low-resource settings
should be ASSURED; affordable, sensitive, specific, user-friendly, rapid and robust, equipment
free, and deliverable to end users (Mabey et al. 2004; Yager et al. 2008). For example,
immunoassay-based strips fit the ASSURED model albeit with limited sensitivity and specificity.
On the other hand, nucleic-acid based assays provide excellent sensitivity and specificity but
often are non-affordable, slow, and require high power. Miniaturized instrument-free LAMP
systems have huge potential for low-cost POC diagnostics.
2.7.3 Simple sample processing in LAMP reactions
Direct gene amplification from bacterial cells would save the time and cost of sample
processing. It will also simplify the device instrumentation and power consumption. Studies have
shown that LAMP assays could be performed from eukaryotic/prokaryotic cells, without getting
affected by cell debris. For example, Iwasaki et al. developed a LAMP assay for genotyping
single nucleotide polymorphism in drug-responsive human cytochrome gene directly from whole
o

blood by heating it at 95 C for 5 min (Iwasaki et al. 2003). LAMP assay was compared against
the gold standard PCR-restriction fragment length polymorphism analysis. Despite the presence
of PCR inhibitory substances such as heme and immunogluobulin G in blood, LAMP assays
were successful in detecting single nucleotide polymorphisms in the presence of cell debris only
in 30 min. On the other hand, PCR-restriction fragment length polymorphism analysis needed
the heat treatment and purification of genomic DNA. Poon et al. demonstrated LAMP assay for
o

detecting Plasmodium falciparum by directly heating the blood at 99 C for 10 min.(Poon et al.
2006) Approximately 100 samples were tested with LAMP and compared against PCR and
microscopic count. PCR was not possible without the purification of genomic DNA. Among 13

69

heat treated blood samples (positive for Plasmodium falciparum) which were also positive by
LAMP, 7 showed negative results by PCR. Clinical sensitivity and specificity of LAMP assay
were 95% and 99% respectively. Also the detection limit of LAMP assay was 6 parasites/ µL,
which was 10-fold higher than the microscopic count method. The same group developed a
reverse-transcription LAMP assay for detecting avian influenza (H5N1) virus directly from viral
o

culture (Jayawardena et al. 2007). Detection sensitivity of heat treated sample (99 C for 10 min)
in LAMP assay was found to be 9.6 PFU/ reaction, which was approximately 1000-fold less
sensitive than assay which used purified RNA. As DNA-based LAMP assays have not shown
any effect of sample heat treatment on the detection sensitivity, this effect might be due to the
sensitivity of RNA towards the heat. In another study, LAMP assay was also performed directly
from oilseed rape plant by mechanically disrupting the tissue (Lee et al. 2009). In all these
o

o

examples, LAMP assays required either prior heating of the sample between 95 C to 99 C (310 min) or a mechanical step for cell lysis before mixing the lysate in the reaction solution.
Additionally, these cell-based LAMP assays were performed in 25 mL PCR vials and not yet
translated on microchips. Due to the requirement of prior heating of sample for cell lysis,
microchip-based LAMP would have an extra manual step. We have found that real-time and
quantitative microLAMP assays could be performed by directly incubating the Gram-negative
o

and Gram-positive bacterial cells in the reaction mixture at 63 C (unpublished). Direct gene
amplification from bacterial cell at the incubation temperature would not require any prior
heating or upstream sample processing step, which would further simplify microLAMP system
instrumentation. Direct cell-based amplification would be highly useful for translating LAMP
assays on microchips with sample-in-answer-out capability.

70

2.8 Conclusions
POC genetic diagnostics critically depends on miniaturization of amplification, detection,
and sequencing techniques. Isothermal amplification techniques, which are still emerging, have a
better potential for POC diagnostics. Isothermal amplification techniques (e.g., loop-mediated
isothermal amplification) are promising for the low-cost and integrated nucleic acid analysis.
Here, parameters that effect the nucleic acid amplification time in microPCR systems are
quantitatively evaluated and analyzed. Till date, the fastest nucleic acid amplification for static
chamber and flow-through micro-PCR system is reported in 5.67 min and 90 s, respectively. For
static chamber microPCR systems, application of integrated thin film heater and highly
conductive microchip lowers the thermal mass, reduce the thermal transition time, increase
heating/cooling rates, and reduce the nucleic acid amplification time. Amplification time for
static chamber microPCR systems can be further reduced (e.g., to 4 min) by applying IR-based
non-contact heaters. However, it would not be a practical option for cost-effective POC
diagnostic devices. Application of high flow-rates and smaller channel cross-sectional areas
reduces the amplification time in flow-through micro-PCR systems. However, to achieve
specificity and reproducibility in microPCR assays, extensive surface treatment and optimization
of flow-rate, channel area, and length are needed. Examples of integrated microPCR systems
show the possibility of POC disease diagnostics. However, they are still expensive to apply for
resource-limited settings. In fact it has been shown that the commercial POC diagnostic systems
based on isothermal assays are considerably cheaper than systems with PCR-based assays.
Examples of instrument-free LAMP systems and simpler sample processing in LAMP assays
have been demonstrated. However, completely integrated LAMP-based nucleic acid
amplification device with on chip sample processing has yet to be achieved. There is possibility

71

to translate microLAMP assays into rapid, sensitive, specific, multiplex, and low-cost diagnostic
platforms. Comparison of microPCR with microLAMP systems shows that the latter is equally
sensitive, specific and rapid in addition to being simpler and less expensive due its isothermal
character. It is essential that nucleic acid amplification steps should be integrated with sample
preparation and detection steps to establish a completely integrated POC infectious disease
diagnostic devices. Almost all the analytical steps of sample processing, amplification, and
detection have reached to a mature stage. The next step should be the integration and automation
of all the functionalities to realize integrated diagnostic systems.

72

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73

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CHAPTER III
3. A CCD-BASED FLUORESCENCE IMAGING SYSTEM FOR REAL-TIME LOOPMEDIATED ISOTHERMAL AMPLIFICATION-BASED RAPID AND SENSITIVE
DETECTION OF WATERBORNE PATHOGENS ON MICROCHIPS

3.1 Introduction
Waterborne bacteria and protozoa that have been commonly associated with disease
outbreaks are Campylobacter jejuni, Escherichia coli O157:H7, Legionella pneumophila,
Shigella flexneri, Salmonella enterica, Vibrio cholerae, Cryptosporidium parvum, and Giardia
intestinalis (Lee 2002; Hoffman et al. 2009). Some of these pathogens such as C. jejuni, E. coli
O157: H7, and S. enterica have also been implicated in outbreaks of food-related illnesses
(Boore et al. 2010). In the United States, 833 documented waterborne disease outbreaks have
occurred between 1971 and 2006 resulting in 577,991 cases of illness and 106 deaths (Craun et
al. 2010). In many cases, these deaths are due to the lack of appropriate diagnosis rather than
availability of effective and economical prevention and treatment options (Yager et al. 2008).
In recent times, isothermal nucleic acid amplification techniques including, helicasedependent amplification (Vincent et al. 2004), recombinase polymerase reaction (Piepenburg et
al. 2006), and loop-mediated isothermal amplification (LAMP) (Notomi et al. 2000) have been
developed for sensitive and specific detection of pathogenic microorganisms. LAMP has
attracted most attention due to its simplicity, ruggedness, and low-cost (Poon et al. 2006). Aside
from its isothermal character, important features of LAMP include, moderate reaction
o

temperature of 63 C, excellent specificity due to the use of 4 to 6 specific primers, need for a
single enzyme, and superior tolerance to substances that typically inhibit PCR (Kaneko et al.

92

2007). Another advantage of LAMP is the generation of high amounts of white magnesium
pyrophosphate precipitate and DNA amplicon (~400 ng /µL) (Mori et al. 2001), allowing a
simpler and inexpensive means for detection. A positive LAMP reaction in 0.2 mL PCR vials is
confirmed by observing the turbidity (Mori et al. 2001) or fluorescence (Chen and Ge 2010),
either in endpoint or real-time formats by using water bath or commercial real-time PCR
instrument (Mori and Notomi 2009). Amplification times of tube-based LAMP assays typically
vary between 30-90 min, depending on the starting DNA template (Seyrig et al. 2011).
Endpoint and real-time LAMP assays employed with turbidity or fluorescence-based
detection schemes have also been translated on microchips for point-of-care diagnostics (Niemz
et al. 2011). Positive endpoint microLAMP is confirmed either visually (Fang 2010) or by
fluorescence microscopy (Matsubara et al. 2002; Lam et al. 2008). Fluorescence (Abe et al.
2011; Liu et al. 2011) or turbidity (Fang et al. 2010) increment during real-time microLAMP is
monitored by photodiodes. MicroLAMP is quantified by measuring the threshold time (Tt) for a
selected cutoff of signal, where the signal is above the background (Fang et al. 2010; Abe et al.
2011). Due to low responsivity (ratio of generated photocurrent to incident photons) and lack of
any exposure control or internal gain features in simple photodiodes (Kostov 2009), reaction
time for real-time microLAMP is generally fixed and depends on the characteristics of LAMP
primers and target DNA, amount of target DNA, and reaction conditions (Kimura et al. 2011). A
same set of LAMP primers under the similar reaction conditions has been evaluated for endpoint
LAMP by gel electrophoresis (En et al. 2008) and real-time turbidity-based microLAMP by
photodiode (Fang et al. 2010). Detection time of approximately 1 h and a sensitivity of 10 fg / µl
were achieved on both formats. Photodetectors such as avalanche photodiodes and silicon

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7

photomultipliers can internally amplify the current up to 10 -fold, drastically enhancing the
responsivity, but they have not yet applied to microLAMP assays (Grigoriev et al. 2007).
On the other hand, increasing the CCD exposure time improves signal-to-background
(Yang et al. 2008) or signal-to-noise ratio (SNR) (Pjescic et al. 2010), enhancing the detection
limit of biochemical assays. Additionally, application of a DNA intercalating dye like Syto-82 to
real-time PCR showed comparatively higher fluorescence signal than the background (Gudnason
et al. 2007). Therefore, increasing the CCD exposure time might increase the SNR and reduce Tt
of Syto-82-based microRTf-LAMP, due to the direct relation between the exposure time and
SNR and inverse relation between SNR and Tt. However, the effect of CCD exposure time on
the SNR or Tt of microRTf-LAMP has not yet been reported. To our knowledge, this is the first
report applying a simple CCD camera for monitoring microRTf-LAMP assays of waterborne
pathogens. Also CCD has wider field-of-view than photodiodes, providing high throughput gene
amplification detection (Dahl et al. 2007).
In this study, a $300 CCD camera, commonly used for astronomical imaging, was
optimized for monitoring RTf-LAMP assays of 6 waterborne pathogens on polymeric
microchips. Also the effect of CCD exposure time on the SNR and Tt of microRTf-LAMP
assays was evaluated. RTf-LAMP assay on microchip required only 2 µL of sample volume and
can be completed within 20 min with sensitivity of a single DNA copy. This low-cost CCDbased fluorescence imaging system can be applied to high throughput microfluidic chips for
rapid and parallel detection of pathogens in low resource settings.

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3.2 Materials and methods
3.2.1 DNA targets
Genomic DNA of S. enterica (700702D), C. parvum (PRA67D), C. jejuni (700819D), L.
pneumophila (33152D), E. coli O157:H7 (BAA460D), and V. cholerae (39315) was obtained
from the American Type Culture Collection (ATCC, Manassas, VA). Before use, dried genomic
DNA was resuspended in nuclease-free sterile water (Fischer Scientific, Pittsburgh, PA). Double
stranded DNA standard (1 kb, 1 µg/µL) was obtained from Invitrogen (Invitrogen Corporation,
Carlsbad, CA).
3.2.2 LAMP primer design
LAMP primers were designed for 11 virulence genes specific to 6 waterborne pathogens
(Table 3.1). For C. jejuni 0414 gene, primers from the literature were used (Yamazaki et al.
2008). Sequences of all primers are provided in the appendix. Prior to primer design, consensus
sequences were generated by aligning multiple gene sequences with Bioedit Sequence
Alignment Editor (Ibis Biosciences, Carlsbad, CA). A set of six specific LAMP primers (F3, B3,
FIP, BIP, LF and LB) were designed for each target by using Primer Explorer V4 (Eiken
Chemicals Co., Tokyo, Japan). Primers were synthesized by Integrated DNA Technologies (IDT,
Coralville, IA).

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Table 3.1 Genes selected for detection of pathogenic microorganisms
Microorganism
Cryptosporidium parvum
Legionella pneumophila
Vibrio cholerae
Campylobacter jejuni
Escherichia coli O157:H7
Salmonella enterica

Gene target(s)
gp60
hsp70
dotA
lepB
ctxA
toxR
0414
cdtA
eae
stx2
invA
phoB

Description
60 kDa glycoprotein
70 kDa heat shock protein
cytoplasmic protein
effector protein
cholera toxin
two-component regulator
oxidoreductase subunit
cytolethal distending toxin
intimin
shiga-toxin 2
invasion protein
response regulator

3.2.3 Microchip fabrication
Microchips with two different designs were fabricated for dsDNA standard dilution series
imaging and monitoring RTf-LAMP assays. Microchips used for dsDNA standard dilution series
imaging consisted of 16 circular wells with 1 mm diameter and 2 µL volume per well.
Microchips for RTf-LAMP had seven V-shaped reaction wells with a volume of 2 µL per well
(Figure 3.1). V-shaped reaction wells provided easy solution dispensing and effective sealing of
microchip. Both types of microchips were fabricated with 100 µm thick ZeonorFilm® (ZF14100; Zeon Chemicals, Louisville, KY) by hot embossing using a sacrificial thermoplastic counter
tool (acrylonitrile butadiene styrene, ABS from K-mac Plastics). Embossing molds were
fabricated via stereolithography (FineLine Prototyping; Raleigh, NC) with the circular and Vshaped features. A sheet of ZeonorFilm® was sandwiched between the ABS plastic and
embossing mold, and preheated to 150 °C in a press (Carver press model 4386; Carver, Wabash,
IN). An embossing pressure of 1000 lbs was applied for 3 min. The system was subsequently

96

cooled to 105 °C while maintaining the embossing pressure. Sample dispensing ports with 750
µm diameter were patterned into the PCR tape (MicroAmp® Optical Adhesive Film; Applied
Biosystems; Foster City, CA) using a commercial-grade knife plotter (CraftROBO model
CC33OL-20; Graphtec). The patterned PCR tape was manually aligned onto the chip, and
pressed at 2000 lbs. with a piece of silicon rubber (70 A, McMaster-Carr, Los Angeles, CA) to
seal the chip. Microchips used for dsDNA standard dilution series imaging were not sealed from
the top.

3.2.4 Experimental setup
Gene amplification and detection system was built by integrating light source, optical
filters, thin film heater, and a CCD camera (Figure 3.1). A 530 nm green LED (05027-PM12,
LED Supply) was used as the excitation light source. The LED was attached to a black anodized
aluminum heatsink (HS13137, LED Supply), optic holder (L2-OH-S35, LED Supply),
collimating lens (L2-OP-025, LED Supply), and an engineered glass diffuser (ED1-C50,
Thorlabs, Newton, NJ) to achieve widely distributed homogenous light. The LED was driven at
700 mA with a power supply (HY5003, Power Supply Depot, Lake Park, FL). The excitation
light was filtered through a 534 ± 20 nm bandpass filter (FF01-534/20-25, Semrock, Rochester,
NY). The emission light was filtered through a 572 ± 20 nm bandpass filter (FF01-572/28-25,
Semrock) connected to the CCD camera. Imaging of microchips for dsDNA standard dilution
series and RTf-LAMP assays was done by a 16 bit, 0.25 megapixel monochrome CCD camera
(MEADE DSI Pro, Irvine, CA) equipped with a 16 mm relay lens (15774, Deal Extreme). DSI
Pro CCD is an interline-scan monochromatic camera commonly applied for astronomical

97

imaging (Gottscheber and Dech 2010). It is equipped with AutoStar Envisage imaging software
with manual control over the gain (0-100), offset (0-100), and exposure time (0.1 ms-1 h). Offset
and gain of the CCD camera were optimized as shown in the Figure 3.2. A gain of 100 and an
offset of 60 were selected for imaging, as this combination provided the most optimum contrast
in the images.

Figure 3.1 Photograph of the experimental setup consisting of an LED attached with excitation
filter (534 ± 20 nm) for illumination, a thin film heater and thermocouple for temperature
control, and a monochromatic CCD camera with emission filter (572 ± 20 nm) for imaging. The
inset shows the microchip with seven V-shaped reaction wells used for RTf-LAMP reactions.
This system was placed in the dark during imaging to avoid any ambient light

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The temperature control system consisted of a thermocouple (5SRTC-TT-T-40-36,
Omega), a pulse width modulator IC (DRV102T, Texas Instruments), a USB DAQs (USB-DAQ
6009, National Instruments), a thin film heater (Electro-Flex Heat Inc.), and a LabVIEW VI
developed in-house. To control the temperature of the microchip at 63

°C,

a feedback

mechanism involving (PID) and pulse width modulation (PWM) was implemented. Signals from
the thermocouple were amplified and conditioned using a LabView thermocouple input module.
The analog output generated from the PID controller was used to adjust the duty cycle of the
PWM driver. The component cost of this experimental setup was approximately $1000.
3.2.5 DNA dilution series and real-time LAMP
Serial dilutions (10-fold) of dsDNA standards ranging from 5 ng / µL to 0.005 ng / µL
were prepared in sterile water. dsDNA standard dilutions were mixed with 2 µM Syto-82.
Negative control had only water and 2 µM of Syto-82 (Invitrogen Corporation, Carlsbad, CA).
Approximately, 2 µL of dsDNA standard dilution series with 10 ng, 1 ng, 0.1 ng, 0.01 ng, and
water were placed on the microchip with circular wells. All experiments were performed in
triplicate.
Reaction mixtures for LAMP contained 1.6 μM each of FIP and BIP primer, 0.200 µM
each of F3 and B3 primer, 0.8 µM each of LF and LB primer, 0.8 M betaine (Sigma, St Louis,
MO), 1.4 mM of each dNTP (Invitrogen Corporation, Carlsbad, CA), 20 mM Tris-HCl (pH 8.8),
10 mM (NH4)2SO4, 10 mM KCl, 8 mM MgSO4, 8 mM Triton X-100, 0.64 units/µL of large
fragment of Bst DNA polymerase (New England Biolabs Inc., Ipswich, MA), 2 µM of Syto-82,
5

and 10 genomic copies of target DNA. To evaluate the dynamics of RTf-LAMP and establish

99

standard curves for quantitative analysis, 10-fold serially diluted genomic DNA of C. jejuni
5

ranging from 10 to 1 copy was prepared. LAMP reaction mixtures without any DNA template
were also prepared to serve as negative controls. Positive and negative LAMP controls with the
volume of 25 µL were loaded in 0.2 mL PCR tubes (VWR International, West Chester, PA) and
incubated at 63 C for 45 min on a commercial real-time PCR instrument (Chromo4

TM

, Bio-

Rad Laboratories, Hercules, CA). Fluorescence signals of the RTf-LAMP assays were measured
every minute in channel 2 (excitation wavelength range from 500-535 nm and emission
wavelength range from 560-580 nm). For RTf-LAMP on microchip, 2 µL of LAMP solutions
was dispensed in each of the 6 wells using a pipettor, manually sealed with a PCR tape, and
o

placed on the experimental setup for incubation at 63 C. All experiments were performed in
triplicate.
3.2.6 Fluorescence measurement, extraction, and analysis
TM

Opticon

software (Bio-Rad Laboratories, Hercules, CA) was used to measure and

extract raw fluorescence signals of RTf-LAMP assays on the real-time PCR instrument. For the
florescence imagining system, CCD camera imaged the microchips and saved the images in
flexible image transport system format. For dsDNA standard dilution series, images of the
microchip (with circular wells) were recorded for increasing exposure times from 1 ms to 60 s.
For microRTf-LAMP assay, time-lapse images of microchip (with V-shaped wells) were
recorded every 30 s at fixed exposure time. Imaging was stopped when the signals from the
positive wells reached to saturation value of 65,535 a. u. CCD images were opened in Image J

100

(National

Institutes

of

Health,

Bethesda,

MD)

and

a

microarray

profile

plugin

(http://www.optinav.com/imagej.html) was applied to extract fluorescence signals from all the
wells. For dsDNA standard dilution series, average fluorescence signals from the triplicates
were plotted with respected to exposure time. To calculate SNR values of RTf-LAMP assays on
real-time PCR instrument and microchip, raw signal intensities were baseline corrected by
subtracting the average fluorescence signals obtained in the first 3 min and dividing by their
standard deviation. Tt was then defined as the time at which the SNR reached an arbitrary cut-off
of 10 or higher. SNR and the corresponding Tt values were calculated for individual LAMP
reaction of triplicates and an average value of Tt is reported here. Fluorescence signals were
normalized by dividing the all the signals with the maximum signal value. Average of the
normalized fluorescence signals from the triplicates was plotted with respect to amplification
time. Data analysis and plotting was done with Microsoft Excel (Microsoft, Redmond, WA).
3.3 Results and discussion
3.3.1 CCD optimization
To observe the effect of offset and gain on the fluorescence signal collected by CCD,
COP microchip containing 10 ng dsDNA standard stained with 2 µM Syto-82 was imaged by the
CCD camera with increasing offset and gain at a fixed exposure time of 1 s. An increase in the
fluorescence signals was observed with increasing gain and offset values (Figure 3.2). Initially
the gain and exposure time was fixed to 100 and 1 s respectively and the offset was varied from
0 to 100 then the offset and exposure time was fixed to 60 and 1 s respectively and the gain was
varied from 0 to 100. For the first condition, fluorescence signal was saturated to 65,535 a. u. at a
gain of 100 and an offset of 80. For the second condition, maximum fluorescence signal of

101

58,283 a. u. was observed at a gain of 100 and an offset of 60. However, for the same gain and
offset of 100 and 60, the fluorescence signal was 62,208 a. u. in the first condition. As the
circular well microchip used for DNA imaging was not sealed, solution evaporation might have
increased the solution concentration, resulting in higher fluorescence signal in the first condition.
A gain of 100 and an offset of 60 were selected for further CCD imaging, as this combination
provided the most optimum contrast in the images.

Figure 3.2 Optimization of offset and gain for the CCD camera. Plot of average fluorescence
signal from 10 ng dsDNA standard stained with 2 µM Syto-82 with respect to increasing offset
and gain at a fixed exposure time of 1 s.

3.3.2 Effect of CCD exposure time on the fluorescence signals of dsDNA standard dilution
series
The effect of CCD exposure time on the fluorescence signals from 10 ng, 1 ng, 0.1 ng,
0.01 ng, and 0 ng dsDNA standard stained with 2 µM Syto-82 were evaluated by imaging the
102

microchip with increasing exposure times from 1 ms to 60 s. Images of dsDNA standard dilution
series on microchip with increasing exposure times are shown in Figure 3.3.

Figure 3.3 CCD images of 16 well microchip with dsDNA standard dilution series (10 ng, 1 ng,
0.1 ng, 0.01 ng, and negative control) stained with 2 µM Syto-82 and with increasing exposure
times

The fluorescence signals from all the DNA concentrations and the negative controls
increased with increasing exposure times and finally reached to saturation (Figure 3.4). From an
exposure time between 1 ms to 0.1 s, fluorescence signals varied between 4,029 a. u. to 4,395 a.
u. (background) and showed an exponential rise from exposure time between 0.1 s to 10 s. After
10 s of exposure time, the fluorescence signals reached the saturation value of 65,535 a. u.
However, the fluorescence signals from 10 ng and 1 ng DNA showed an increase from 0.01 s
exposure time. Also the fluorescence signals from 10 ng and 1 ng DNA were relatively higher
than the signals from the negative control for the exposure time between 0.5 s to 5 s. For

103

example, fluorescence signals from 10 ng and 1 ng DNA were approximately 8-fold and 3-fold
higher than negative control at 0.5 s of CCD exposure time and 6.5-fold and 5-fold higher than
the negative control at 5 s of CCD exposure time, respectively. The lower fluorescence signals
from the negative control in comparison to 10 ng and 1 ng DNA standard were due to the low
autofluorescence of Syto-82 dye. Syto-82 has a substituted unsymmetrical cyanine structure,
providing very low autofluorescence in unbound state due to flexibility of its structure in
aqueous environment. This cyanine structure becomes rigid in the presence of nucleic acids,
providing high fluorescence signals (Tarnok 2008; Deligeorgiev et al. 2009). (Gudnason et al.
2007) demonstrated a 50-fold lower detection limit with 2 µM Syto-82 in real-time PCR assay,
when compared with 14 other DNA intercalating dyes. Fluorescence signals from 0.1 ng and
0.01 ng DNA were nearly similar to the fluorescence signals from the negative control for the
applied CCD exposure times. This effect might be due to the lower amount of dsDNA standard
available to bind with Syto-82 dye, producing the fluorescence signals similar to the negative
control. We hypothesized that applying the exposure time effect to microRTf-LAMP might lead
to faster increment in SNR due to the collection of relatively higher fluorescence signals from
the positive controls than the background, reducing the Tt of assay.

104

Figure 3.4 Plot of average fluorescence signal with respect to increasing exposure times for the
dilution series of dsDNA standard (10 ng, 1 ng, 0.1 ng, 0.01 ng, and 0 ng) stained with 2 µM
Syto-82

3.3.3 Effect of CCD exposure time on SNR and Tt values of microRTf-LAMP assays
The effect of CCD exposure time on the SNR and Tt were evaluated for two different
5

LAMP assays for V. cholerae toxR gene and C. parvum gp60 gene (10 gene copies) for the
exposure time of 1 s, 3 s, and 5 s. CCD exposure time was limited to 5 s, as applying an exposure
time higher than 5 s either saturated the signal in the positive wells or provided a high
background. The SNR and Tt values of microRTf-LAMP assays at these exposure times were
also compared with a real-time PCR instrument. A systematic increment in the SNR and
reduction in the Tt values of microRTf-LAMP assays were obtained with increasing exposure

105

time as compared to the real-time PCR instrument (Table 3.2 and Figures 3.5 and 3.6). Negative
controls did not show any increase in the fluorescence signal.
5

Table 3.2 Comparison of average SNR and Tt values of RTf-LAMP for 10 DNA copies of V.
cholera toxR gene and C. parvum gp60 gene on microchips (1 s, 3 s, and 5 s CCD exposure
time) and the real-time PCR instrument. Standard deviations are the mean of Tt values from
triplicates
Microorganism/gene

V. cholera
toxR
C. parvum
gp60

SNR, Tt (min)
CCD system

SNR, Tt (min)
CCD system

SNR, Tt (min)
CCD system

1 s exposure
21.8, 5.8 ± 0.4

3 s exposure
14.8, 5 ± 0.0

5 s exposure
70.8, 5 ± 0.0

SNR, Tt (min)
TM
Chromo4
28.3, 7.7 ± 0.6

12.1, 18.7 ± 0.3

12.6, 15 ± 0.0

206, 14.5 ± 0.0

16.3, 24.3± 1.2

For microRTf-LAMP assay of V. cholerae toxR gene, SNR and Tt values for 1 s, 3 s, and
5 s CCD exposure times were 21.8 and 5.8 min, 14.8 and 5 min, and 70.8 and 5 min respectively.
Tt at 3 s and 5 s of CCD exposure time was same as 5 min. However, SNR at 5 s of exposure
time was approximately 5-fold higher than SNR at 3 s of exposure time. SNR and Tt for the
same assay on the real-time PCR instrument were 28.3 and 7.7, respectively. Moreover, Tt
values for the microRTf-LAMP assay with all the applied CCD exposure times were lower than
the Tt obtained on the real-time PCR instrument (Figure 3.5). Time lapse images of microchips
during the amplification of V. cholerae toxR gene at 1s, 3s, and 5 s of exposure time showed that
higher fluorescence signals were achieved faster at higher exposure time (Figure 3.7).

106

5

Figure 3.5 RTf-LAMP curves for 10 DNA copies of V. cholera toxR gene on the microchips (1
s, 3 s, and 5 s CCD exposure time) and real-time PCR instrument

5

Figure 3.6 RTf-LAMP curves for 10 DNA copies of C. parvum gp60 gene on the microchips (1
s, 3 s, and 5 s CCD exposure time) and real-time PCR instrument

107

For microRTf-LAMP assay of C. parvum gp60 gene, SNR and Tt values for 1 s, 3 s, and
5 s CCD exposure time were 12.1 and 18.7 min, 12.6 and 15 min, 206 and 14.5 min,
respectively. A difference in Tt value of 0.5 min was observed for microRTf-LAMP assay with 3
s and 5 s of CCD exposure time. However, SNR at 5 s of CCD exposure time was 16-fold
higher than the SNR at 3 s of CCD exposure time. SNR and Tt values for the same assay on the
real-time PCR instrument were 16.3 and 24.3 min, respectively. A 9.8 min difference in Tt value
of microRTf-LAMP with 5 s of CCD exposure time and real-time PCR instrument was observed
(Figure 3.6). It was recently demonstrated that the reaction time is highly dependent on the
physical properties of LAMP primers such as GC content, melting temperature, and free energy
of hybridization (Kimura et al. 2011). As the same LAMP reaction conditions had been applied,
the difference in Tt values of these two LAMP assays (V. cholerae toxR gene and C. parvum
gp60 gene) was due to the difference in the characteristics of LAMP primers and target DNA.
Time lapse images of microchips during the amplification of V. cholerae toxR gene at 1s,
3s, and 5 s of exposure time showed that higher fluorescence signals were achieved faster at
higher exposure time (Figure 3.7).

108

Figure 3.7 Time lapse CCD images of microRTf-LAMP assay for V. cholera toxR gene with
CCD exposure times of 1 s, 3 s, and 5 s. Fluorescence signals increase with the increasing
exposure times

Amplicon yield of a 0 min turbidity-based LAMP with 6000 starting DNA copies was
5

shown to be approximately 400 ng/µl (Mori et al. 2001). Our microRTf-LAMP assays had 10

starting DNA copies, reaction volume of 2 µl, and total amplification time of 20 min for 5 s CCD
exposure time. If we assume similar reaction kinetics, our microRTf-LAMP assays should
generate more than 267 ng amplicons in 2 µl reaction volume in 20 min. That would produce
approximately 70 ng amplicons in 5 min (minimum Tt value of microRTf-LAMP), which should
provide higher fluorescence from positive microRTf-LAMP relative to the background at higher
exposure times. This effect was similar to the increase in the fluorescence signals from 10 ng
and 1 ng dsDNA standard in comparison to the negative control by increasing the CCD exposure
times (Figure 3.4). Other studies have observed the increment in signal-to-background ratio from
chemiluminescence immunoassay with increasing CCD exposure times (Yang et al. 2008).

109

Similarly, SNR values of positive microRTf-LAMP assays showed consistent increment (relative
to the background) with increasing exposure times, reducing the Tt values of these assays.
However, this effect could not be expected from the fluorescence LAMP on real-time PCR
instrument, which employs photodiode detector.

3.3.4 MicroRTf-LAMP assays of waterborne pathogens
CCD-based imaging system was validated for 12 virulent genes of 6 waterborne
pathogens using Syto-82 dye and real time fluorescence LAMP on V-shaped microchips. Major
waterborne pathogenic bacteria including, S. enterica (invA and phoB gene), C. jejuni (0414 and
cdtA gene), L. pneumophila (dotA and lepB gene), E. coli O157:H7 (stx2 and eae gene), and V.
cholerae (toxR and ctxA gene), and a protozoan, C. parvum (hsp70 and gp60 gene) were
selected from the list of Centers for Disease Control and Prevention (Lee 2002). MicroRTf5

LAMP assays for 12 virulent genes of waterborne pathogens (10 gene copies) were performed
at 5 s of CCD exposure time and compared with the real-time fluorescence LAMP on a real-time
PCR instrument (data not shown). A reduction in Tt values ranging from 2.7 min to 9.8 min was
achieved for microRTf-LAMP assays in comparison to real-time PCR instrument (Table 3.3).
LAMP primer sets, which were slower (higher Tt) on real-time PCR instrument showed higher
reduction in the Tt values on microchip than the LAMP primers which were relatively faster
(lower Tt) on the real-time PCR instrument For example, LAMP primer set for C. parvum gp60
gene showed the slowest amplification on the real-time PCR instrument with Tt of 24.3 min,
which was reduced by 9.8 min on the microchip assay. LAMP primer set for V. cholerae toxR
gene was the fastest on the real-time PCR instrument with an amplification time of 7.7 min,

110

which was reduced by 2.7 min on microchip assay. Moreover, standard deviations in the Tt
values of microRTf-LAMP assays were either same or lower than the real-time PCR instrument.

5

Table 3.3 Comparison of average Tt values of RTf-LAMP for 10 DNA copies of 6 waterborne
pathogens (2 genes for each) on the microchips with 5 s CCD exposure and the real-time PCR
instrument. Standard deviations are the mean of Tt values from triplicates
Microorganism

Gene

S. enterica

invA

Tt (Chromo4
(min)
11 ± 0.0

S. enterica

phoB

C. parvum

TM

)

Tt (CCD System)
(min)

Δ t (min)

6.5 ± 0.0

4.5

14.7 ± 0.6

10 ± 0.0

4.7

hsp70

15.3 ± 0.6

11 ± 0.0

4.3

C. parvum

gp60

24.3 ± 1.2

14.5 ± 0.0

9.8

C. jejuni

0414

16.7 ± 0.6

7.8 ± 0.6

8.9

C. jejuni

cdtA

16 ± 0.0

9.3 ± 0.3

6.7

L. pneumophila

dotA

10 ± 0.0

5.5 ± 0.5

4.5

L. pneumophila

lebB

11.6 ± 0.6

6.5 ± 0.0

5.5

E. coli O157:H7

stx2

11 ± 0.0

4.8 ± 0.3

6.2

E. coli O157:H7

eae

18.3 ± 0.6

13.3 ± 0.3

5.0

V. cholerae

toxR

7.7 ± 0.6

5 ± 0.0

2.7

V. cholerae

ctxA

17.3 ± 0.6

10 ± 0.0

7.3

LAMP primer set for C. jejuni 0414 gene was evaluated for the dilution series of 14, 1.4,
and 0.14 DNA copies, by a real-time turbidimeter (Yamazaki et al. 2008). Average Tt values for
real-time fluorescence LAMP assays in our study were similar to their reported average values of
Tt for real-time turbidity (Yamazaki et al. 2008). However, there was a high deviation in their Tt
values between the replicates (not discussed). Similarly, (Chen and Ge 2010) observed the
variations in the Tt values of the replicates of real-time turbidity-based LAMP, when compared
with real-time fluorescence-based LAMP for the same target of V. parahaemolyticus toxR gene.
Other studies have also reported the real-time fluorescence LAMP assays to be faster, sensitive,
111

and more reproducible than real-time turbidity-based LAMP (Aoi et al. 2006; Chen and Ge
2010). We also observed very low standard deviations in the Tt values of real-time fluorescence
LAMP assays on microchips and real-time PCR instrument. Lee et al. (2008) reported an optical
system equipped with a photodiode for monitoring real-time turbidity-based microLAMP assay.
A slight variation in the intensity of excitation light or its alignment to the detector was able to
affect the turbidity value due to the additional scattering of photons (Lee et al. 2008). However,
fluorescence signals from CCD-based microchip assays such as microPCR (Dahl et al. 2007) or
microRTf-LAMP (in this study) are not affected by the slight sample misalignment.

3.3.5 MicroRTf-LAMP assay for quantitative analysis of C. jejuni
To establish standard curves for quantitative analysis, real-time fluorescence LAMP
assays targeting 0414 gene were performed on microchips with 5 s CCD exposure time and on a
5

real-time PCR instrument with 10-fold serial dilutions of C. jejuni DNA ranging from 10 to 1
copy. The standard curves were established by plotting the Tt values versus log of the number of
genomic DNA copy used in RTf-LAMP assays (Figure 3.8). Both the microRTf-LAMP and RTfLAMP assay on commercial real-time PCR instrument were sensitive to a single DNA copy.

112

Figure 3.8 Standard curves for the C. jejuni 0414 gene amplification on the microchips at 5 s of
CCD exposure time and real-time PCR instrument

Yamazaki et al. (2008) had also reported a single copy level sensitivity for this primer
set. The correlation coefficients of the log linear regression plots between Tt values and DNA
copy numbers for microRTf-LAMP and real-time PCR instrument were the same, 0.99.
However, microRTf-LAMP assay was approximately twice faster than the RTf-LAMP assay on
the real-time PCR instrument. The result indicates that microRTf-LAMP enables the
reproducible and rapid quantification of DNA.

113

3.4 Conclusions

In this study, we demonstrated that rapid RTf-LAMP assays can be performed on COP
microchips in comparison to a commercial real-time PCR instrument. Increasing the CCD
exposure time up to 5 s, increased the SNR by 8-fold and reduced the Tt values from 2.7 min to
5

9.8 min for 10 starting DNA copies of microRTf-LAMP assays. Additionally, microRTf-LAMP
assay allowed the analysis of low sample volume of 2 μL in less than 20 min with a detection
limit of single DNA copy. Due to requirements of single temperature incubation, microRTfLAMP assays have the potential to be translated into rapid and low cost battery operated
molecular diagnostic devices, suited to applications in resource-limited settings.

114

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115

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CHAPTER IV
4. DIGITAL LAMP FOR RAPID AND SENSITIVE GENETIC AMPLIFICATION OF
GRAM-NEGATIVE AND GRAM-POSITIVE BACTERIA WITH NO SAMPLE
PROCESSING
4.1 Introduction
Sample preparation is a key aspect in the implementation sample-in-answer-out gene
analysis systems. Sample preparation methods to recover nucleic acids entails several steps
depending on the sample being analyzed, but often includes cell enrichment, cell lysis and
nucleic acid purification (Mariella 2008; Kim et al. 2009; Price et al. 2009).
Enrichment/concentration of target cells becomes extremely important in the sensitive diagnosis
of infectious disease agents due to their low concentrations in high amount of background (e.g.,
9

200 human immunodeficiency virus particles in 1 mL of blood containing 5×10 red blood
cells). Sample preparation techniques such as filtration (Mach and Di Carlo 2010), immunoseparation (Dharmasiri et al. 2010), and dielectrophoresis (Hu et al. 2005) for cell enrichment,
and biochemical, thermal, electrical, and mechanical methods for cell lysis (Brown and Audet
2008), and silica-based solid-phase methods for nucleic acid separation and purification (Price et
al. 2009) have been translated on microfluidic chips. Microfluidic chips (for sample preparation)
are being integrated with nucleic acid amplification and detection modules for gene analysis
systems with sample-in-answer-out capabilities (Easley et al. 2006). However, microfluidicbased sample preparation modules are limited by low nucleic acid extraction efficiency (50-80%)
and high processing time (~1 h), affecting the sensitivity and total detection time of gene analysis
systems (Grodzinski et al. 2003). As such, development of rapid gene analysis systems would
require either the elimination of sample preparatory steps or improvements in their protocols.

120

Microfluidic PCR systems are known to carry out DNA amplification directly from
bacterial cells (Kaigala et al. 2006; Atrazhev et al. 2010; Manage et al. 2011) by using
polymerases which are less prone to inhibition (Kermekchiev et al. 2009; Matheson et al. 2010;
Zhang et al. 2010). Taq polymerase, the enzyme commonly used to amplify nucleic acids in PCR
is inhibited by a variety of compounds such as cell debris, hemoglobin, and humic acid (Abu AlSoud and Radstrom 1998). Therefore, microfluidic PCR systems are often coupled with
dedicated cell lysis and nucleic acid purification steps to remove cell debris and compounds that
could inhibit DNA amplification (Easley et al. 2006; Witek et al. 2006; Park et al. 2011).
Loop-mediated isothermal amplification (LAMP) is a robust molecular method initially
reported in 2000 and extensively been applied for the DNA/RNA-based amplification of a
variety of pathogens (Notomi et al. 2000; Seyrig et al. 2011; Tourlousse et al. 2011). LAMP is
highly sensitive, specific, and rapid, and most importantly, it is not affected by the presence of
cell debris and compounds which typically inhibit PCR (Masaomi et al. 2003; Poon et al. 2006).
o

LAMP is a promising method for nucleic acid amplification from bacterial cells at 63 C.
LAMP with bacterial cells (in conventional PCR vials) is performed with a separate heating of
bacterial cells and mixing of cell lysate in the reaction solution. Direct heating of reaction
o

solution above 80 C deactivates Bst polymerase (Matovu et al. 2010). In situ LAMP is a
modification of LAMP for DNA amplification directly from bacterial cells (Maruyama et al.
2003; Ye et al. 2011). In situ LAMP requires prior fixation and permeabilization of bacterial
o

cells on a solid surface followed by LAMP reaction at 63 C and fluorescence imaging of cells.
Nucleic acids are amplified inside the cells by the inclusion of LAMP reaction components in the
permeabilized cells. Due to the steps involved in sample processing, in situ LAMP is performed

121

in 3-4 h. DNA-based digital Polymerase chain reaction (Pohl and Shih 2004) and digital
recombinase polymerase amplification (Shen et al. 2011) are performed for quantifying low
concentrations of nucleic acids. Digital LAMP from bacterial cells would be more effective in
target qualification, due to the elimination of upstream sample processing and the loss of nucleic
acids associated with the processing.
To our knowledge, there is no report of a simple, sensitive, and rapid method for nucleic
o

acid amplification directly from bacterial cells at a moderate temperature of 63 C. Direct cellbased microfluidic LAMP would reduce the instrument complexity as compared to microfluidic
PCR due to the isothermal nature of LAMP reaction. Direct cell-based LAMP without any
sample processing is yet to be translated on microfluidic platform. Here we report an isothermal
method for gene amplification directly from gram-negative (Escherichia coli) and gram positive
o

(Enterococcus faecalis) bacterial cells at 63 C with most probable number (MPN) approach
leading to the concept of “digital LAMP”. Real-time fluorescence loop-mediated isothermal
amplification (RTf-LAMP) assays were performed with bacterial cells, heat treated bacterial
o

cells (95 C for 5 min), and their purified genomic DNA templates on a commercial real-time
PCR instrument. Gene amplification from viable bacterial cells was confirmed by performing
propidium monoazide (PMA)-RTf-LAMP assay with E. faecalis cells. Single cell-based RTfLAMP assays were also translated on polymeric microchips (microRTf-LAMP) and monitored
by a fluorescence imaging system, previously developed by our group (Ahmad et al. 2011).
Single cell-level genetic amplification was performed by endpoint microRTf-LAMP assays (10
5

5

cells to 0.01 cell) as digital LAMP. RTf-LAMP assays (from 10 cells to single cell) of E. coli

122

(uidA gene) and E. faecalis (gelE gene) were successfully performed by directly incubating the
o

cells at 63 C without any prior processing. Threshold time values of RTf-LAMP assays of
o

bacterial cells, heat treated bacterial cells (95 C for 5 min), and their purified genomic DNA
templates were similar, implying that DNA could be amplified either by lysis or permeabilization
o

of bacterial cells at 63 C. PMA-LAMP results confirmed that DNA amplification was achieved
with viable bacterial cells and not with the dead cells or extracellular DNA. MicroRTf-LAMP
assays provided signal-to-noise (SNR) ratio between 500- to 800, sensitivity of single cell, and
detection time of approximately 20 min. MPN values from digital LAMP assays confirmed the
genetic amplification were from single cell. Direct gene amplification from bacterial cells at 63
o

C would eliminate the requirement of complex sample preparation steps, facilitating the

development of sensitive, rapid, low-cost, and integrated point-of-care diagnostic systems.
4.2 Materials and methods
4.2.1 Bacterial cell and genomic DNA
E. faecalis (19433) cells were obtained from ATCC. E. coli K12 strain C3000 cells were
graciously provided by Professor Joan B. Rose, Water Quality, Environmental, and Molecular
Microbiology Laboratory at Michigan State University. Frozen stocks of E. coli and E. faecalis
were resuspended in sterile Trypticase Soy Broth (TSB, 211768, BD, Sparks, MD). To revive
o

the cells, these tubes were incubated at 37 C in a shaker-incubator (New Brunswick Scientific,
Edison, New Jersey) at 150 rpm for 12 h. Next, cultures were inoculated into 25 mL of TSB and
o

incubated at 37 C for 8-12 h to achieve mid-log phase. Bacterial cultures were placed overnight

123

o

in the refrigerator at 4 C. Cultures were serially diluted in TSB and colony forming units of E.
coli and E. faecalis were enumerated by plating 0.1 mL of dilutions of cultures on Trypticase
o

Soy Agar (TSA, 211043, BD, Sparks, MD) plates. TSA plates were incubated at 37 C for 12 h.
8

Cell concentrations were approximately 1×10 per mL for both the bacterial species. Strict
aseptic techniques were followed throughout the experiments. Genomic DNA of E. coli strain K12 (700926D) and E. faecalis strain V583 (700802D) was obtained from American Type Culture
Collection (ATCC, Manassas, VA). Before use, dried genomic DNA was resuspended in
nuclease-free sterile water (Fischer Scientific, Pittsburgh, PA).
4.2.2 LAMP primer design
LAMP primers were designed for β-D glucuronidase (uidA) and glutamate decarboxylase
(gadA) genes of E. coli K-12 strain C3000, and gelatinase (gelE) gene of E. faecalis strain V583
respectively. Sequences of all primers are shown in the appendix. A set of six specific LAMP
primers (F3, B3, FIP, BIP, LF and LB) were designed for each target by using Primer Explorer
V4 (Eiken Chemicals Co., Tokyo, Japan). Primers were synthesized by Integrated DNA
Technologies (IDT, Coralville, IA).
4.2.3 Microchip fabrication
Microchips with circular wells were fabricated with approximately 127 µm thick
polyester film (8567K52, McMaster-Carr, Los Angeles, CA). A perforated steel sheet (10 cm ×
12 cm) was gifted by Diamond Manufacturing Company, Wyoming, PA. The perforated sheet
had 6600 holes, arranged in a straight line pattern, each having a diameter of 1 mm and hole-hole
distance of 0.5 mm. Polyester was selected due to its low cost ($22 for 100 cm ×300 cm sheet)
124

and biocompatibility to LAMP reaction. A sheet of polyester was sandwiched between a
perforated steel plate (Perforated Metals Plus, Charlotte, NC) and a piece of silicon rubber (70 A,
McMaster-Carr, Los Angeles, CA), and preheated to 220 °C in a press (Carver press model
4386; Carver, Wabash, IN). An embossing pressure of 1000 lbs was applied for 3 min. The
system was subsequently cooled to 105 °C while maintaining the embossing pressure. Microchip
was then extracted from perforated sheet, rinsed with distilled water, and air-dried at 65°C in an
oven. Microchip consisted of multiple circular wells with 1 mm diameter and 2 µL volume per
well.
4.2.4 Experimental setup
Gene amplification and detection system was built by integrating light source, optical
filters, thin film heater, and a CCD camera. Details of the fluorescence imaging system are
provided in a previous publication (Ahmad et al. 2011). This system was slightly modified to
obtain homogenous light for high throughput imaging. In the modified version, light emitting
diode (LED) (05027-PM12, LED Supply) was replaced by a 525 nm LED array (LIU002;
Thorlabs, Newton, New Jersey). LED array consists of 20 individual LEDs on a printed circuit
board, providing homogeneously distributed light on larger area. Other components such as
excitation filter (FF01-534/20-25, Semrock, Rochester, NY), emission filter (FF01-572/28-25,
Semrock) filter, CCD camera with AutoStar Envisage imaging software (MEADE DSI Pro,
Irvine, CA), and lens (16 mm, 15774, Deal Extreme), thermocouple (5SRTC-TT-T-40-36,
Omega), pulse width modulator IC (DRV102T, Texas Instruments), USB DAQs (USB-DAQ
6009, National Instruments), thin film heater (HK5165R52.3L12, Minco, Minneapolis, MN), and

125

LabVIEW VI (developed in-house) were same as the previously reported system. The
component cost of this experimental setup was approximately $1200.

4.2.5 RTf-LAMP on commercial PCR instrument and on microchip
Reaction solution for RTf-LAMP contained 1.6 μM each of FIP and BIP primer, 0.200
µM each of F3 and B3 primer, 0.8 µM each of LF and LB primer, 0.8 M betaine (Sigma, St
Louis, MO), 1.4 mM of each dNTP (Invitrogen Corporation, Carlsbad, CA), 20 mM Tris-HCl
(pH 8.8), 10 mM (NH4)2SO4, 10 mM KCl, 8 mM MgSO4, 8 mM Triton X-100, 0.64 units/µL of
large fragment of Bst DNA polymerase (New England Biolabs Inc., Ipswich, MA), 0.2%
Pluronic® F-68 (Invitrogen, Carlsbad, CA), 1 mg/ml bovine serum albumin (New England
BioLabs), and 2 µM of Syto-82 (Invitrogen Corporation, Carlsbad, CA). Depending on the type
of RTf-LAMP assay (DNA or cell-based), LAMP solution was mixed with 1 µL of DNA
5

5

solution (10 genomic DNA templates/ µL) or bacterial solution (10 cells/ µL) without any
prior treatment. To examine the effect of heating on cell lysis, cell solutions were separately preo

heated at 95 C for 5 min before mixing in the LAMP reaction solution. In another experiment,
o

LAMP reaction solutions mixed with bacterial cells were preheated at 95 C for 5 min to
examine the effect of heating on Bst polymerase and on the performance of LAMP. To establish
standard curves for quantitative analysis of E. coli and E. faecalis, RTf-LAMP assays with 105

fold serial dilutions of bacterial cells (10 cells -1 cell) were prepared. RTf-LAMP assays
without any DNA template or bacterial cell were also prepared to serve as negative controls.
Positive and negative LAMP controls with the volume of 25 µL were loaded in 0.2 mL PCR

126

tubes (VWR International, West Chester, PA) and incubated at 63

C

for 30-40 min on a

TM

commercial real-time PCR instrument (Chromo4

, Bio-Rad Laboratories, Hercules, CA).

Fluorescence signals of the RTf-LAMP assays were measured every minute in channel 2
(excitation wavelength range from 500-535 nm and emission wavelength range from 560-580
nm). RTf-LAMP assays were performed in triplicate except 6 replicates for single cell dilution.
For single cell RTf-LAMP assays on microchip, 1 µL of E. coli and E. faecalis solution
(1 cell/ µL) was dispensed in each well (n=10) followed by 1 µL of LAMP solution. Wells were
covered with mineral oil (Sigma, St. Louis, MO) during the reaction to avoid solution
o

evaporation. Microchip was placed on the heating stage for incubation at 63 C for 30 min and
real-time imaging.

4.2.6 PMA-RTf-LAMP assays with E. faecalis cells
In order to detect viable E. faecalis in the presence of dead cells and extracellular DNA
templates in the cell solution, RTf-LAMP assays were performed with DNA-modifying PMA
dye. PMA stock solution (20 mM, Biotium Hayward, CA) was diluted to 500 µM in dimethyl
5

sulfoxide. Approximately 9 µL of cell solutions (10 cells/µL) were placed in three PCR tubes
o

(0.2 mL). The solution of the second tube was preheated at 95 C for 5 min to lyse the cells.
Solutions in the first and third tubes were not treated with heat. Approximately 1 µL of 500 µM
PMA solution was mixed in the solutions of first and second PCR tubes to achieve PMA
concentration of 50 µM. These two tubes were incubated in the dark for 20 min followed by

127

exposure with blue light (450 nm LED, LXML-PR01-0500, Ontario, Canada) for 15 min.
Solution in the third tube was neither mixed with PMA nor exposed to the light. For RT f-LAMP
o

assays, 1 µL of each cell solution was mixed with 24 µl LAMP solutions and incubated at 63 C
for 30 min on a commercial PCR instrument. Solutions from these three tubes were streaked on
o

TSA plates and incubated for 12 h at 37 C.

4.2.7 Digital LAMP assays for most probable number calculation
In order to confirm the amplification from the single bacterial cell and to calculate most
probable number (MPN), endpoint microLAMP assays were performed with serial dilutions of
5

E. coli (gadA gene) and E. faecalis (gelE gene) from 10 cells to 0.01 cell for digital LAMP.
Digital LAMP assay for gadA gene of E. coli was designed and tested for single cell-level
sensitivity. Each of the 40 wells was dispensed with 1 µL of E. coli / E. faecalis dilution solution
followed by 1 µL of LAMP solution. Five replicates of digital LAMP assays for each dilution
were performed to match with the 5-tube MPN analysis. Wells were covered with mineral oil
o

and the microchip was incubated at 63 C for 40 min before imaging.

4.2.8 Fluorescence signal analysis
TM

Opticon

software (Bio-Rad Laboratories, Hercules, CA) was used to extract the raw

fluorescence signals of RTf-LAMP assays on the commercial real-time PCR instrument. For
microRTf-LAMP assays, CCD imaged the microchips in time-lapse mode at every 30 s at an
exposure time of 5 s. Imaging was stopped when the signals from the signals from the positive

128

wells reached to a plateau. A gain of 100 and an offset of 60 were used for CCD imaging.
Images were saved in flexible image transport system format. CCD images were opened in
Image J (National Institutes of Health, Bethesda, MD) and a microarray profile plugin
(http://www.optinav.com/imagej.html) was applied to obtain fluorescence signals from the
images. A microarray grid (circular wells with diameter of 1 mm) was used to extract the signals
from the same pixel region of all the time lapse images. To calculate SNR values of RTf-LAMP
assays on the commercial real-time PCR instrument and on microchip, raw fluorescence signals
were baseline corrected by subtracting the average fluorescence signals obtained in the first 3
min and dividing by their standard deviation. Average of the SNR from the triplicates was
plotted with respect to amplification time. Tt was then defined as the time at which the SNR
reached an arbitrary cut-off of 10 or higher. Tt values were calculated for individual LAMP
reaction of triplicates and an average value of Tt is reported here. The error bars in the plots
represented the standard deviations of the mean from three replicate experiments unless
otherwise indicated. Data analysis and plotting was done with Microsoft Excel (Microsoft,
Redmond, WA). Contrast of the images of single cell microRTf-LAMP and endpoint
microLAMP assays were modified for visual clarity. Modification in the image contrast would
not change the fluorescence signal values.
4.3 Results and discussion

4.3.1 RTf-LAMP assays from genomic DNA templates and bacterial cells
5

5

RTf-LAMP assays with 10 genomic DNA copies and 10 bacterial cells of E. coli (uidA
o

gene) and E. faecalis (gelE gene) were performed at 63 C on a commercial PCR instrument. To

129

ensure bacterial cell lysis during the reaction, RTf-LAMP assays were also performed with both
o

the type of bacterial cells, separately preheated at 95 C for 5 min. RTf-LAMP assays were also
o

performed, in which LAMP solution containing bacterial cells were preheated at 95 C for 5 min
to examine the effect of heating on the performance of Bst polymerase. RTf-LAMP assays with
genomic DNA templates and bacterial cells of E. coli (gram-negative) and E. faecalis (gramo

positive) showed the DNA amplification at 63 C (Figure 4.1). DNA amplification was also
achieved from the RTf-LAMP assays of bacterial cells, in which the bacterial cells were
separately pretreated with heat (data not shown). No DNA amplification was achieved in the
RTf-LAMP assays of bacterial cells, in which the LAMP reaction solutions containing bacterial
cells were pretreated with heat (data not shown). Negative controls did not show any increase in
the SNR value during the reaction.
Gram-positive and gram-negative bacteria contain peptidoglycan layer, which is crosslinked with peptide bonds to provide the rigidity to the cell walls (Cabeen and Jacobs-Wagner
2005). Peptidoglycan layer in gram-positive microorganisms is thicker (15-80 nm) and more
cross-linked than the gram-negative bacteria (~10 nm) resulting in more robust cell wall. LAMP
assays have been reported with a number of prokaryotic and eukaryotic cells (Iwasaki et al.
2003). However, in all the cases, prior heating of sample containing bacterial cells at
o

approximately 95 C (3-10 min) for cell lysis and centrifugation to separate bigger cell debris
were performed (Poon et al. 2006; Chen and Ge 2010). To our knowledge, this is the first
demonstration of DNA amplification directly from bacterial cells (without any processing) at 63
o

C.

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5

5

Figure 4.1 RTf-LAMP assays of 10 genomic DNA copies and 10 cells of E. coli (uidA gene)
o
and E. faecalis (gelE gene) at 63 C on a commercial PCR instrument

Average SNR of bacterial cell assays were higher than genomic DNA-based assays. Also
the average SNR from the cell assay of E. faecalis was higher than the cell assay of E. coli. For
example, the highest average SNR from E. coli cell assay and genomic DNA assay were 725 and
352 respectively. The highest average SNR from E. faecalis cell assay and genomic DNA assay
were 1893 and 1275 respectively. The highest average SNR from E. faecalis cell assay was 2.6fold higher than E. coli cell assay. Although Syto-82 is a DNA interacting dye, higher SNR from
bacterial cell assays than genomic DNA-based assays might be due to some interaction of Syto82 dye with the cell membranes in addition to the genomic DNA. For another experiment, E. coli

131

growth was monitored in the presence of 2 µM Syto-82. Syto-82 was not appropriate for cell
growth monitoring, but it showed a 2-fold higher fluorescence signal in the presence of E. coli
cells than in water. Therefore, the interaction of Syto-82 dye with the cell membrane cannot be
completely ruled out. Also higher SNR from E. faecalis cell assay than E. coli cell assay
supports this hypothesis due to the thicker peptidoglycan layer in gram-positive, E. faecalis than
gram-negative, E. coli. Higher standard deviations (SNR) in the cell assays than genomic DNA
assays were observed, which might be due to non-specific interactions of Syto-82 with cell
membranes.
Average Tt values of genomic DNA-based assays were slightly lower than bacterial cell
assays. For example, the average Tt values of E. coli genomic DNA and cell assay were 6 min
(SNR: 19.97) and 7 min (SNR: 13.89) respectively. And the average Tt values of E. faecalis
genomic DNA and cell assay were 6 min (SNR: 161.1) and 7 min (SNR: 49.87) respectively.
o

Also the average Tt values of the bacterial cells, which were preheated at 95 C for 5 min were
exactly the same as the cells without any preheating (date not shown). RT f-LAMP assays
o

containing bacterial cells, which were directly heated at 95 C for 5 min, did not show any
o
amplification. This was due to the fact that direct heating of LAMP reaction mixture at 95 C

deactivated Bst polymerase (Matovu et al. 2010).
o

o

Effect of temperature shock (from 50 C to 70 C) on the membrane permeability of a
gram-negative (E. coli) and a gram-positive (Deinococcus radiodurans) was schematically
studied by using flow cytometry (Baatout et al. 2005). An increase in the cell permeability of
o

both the bacterial was observed at temperatures higher than 45 C. Therefore, in our case,

132

o

incubation of cells at 63 C might have increased the cell permeability leading to the inclusion of
LAMP reaction components in the cells and DNA amplification inside the cells as in situ LAMP
(Maruyama et al. 2003; Ye et al. 2011). The results showed that gene amplification was possible
o

either by lysis or permeabilization of gram-negative and gram-positive bacterial cells at 63 C.
Genomic DNA assays were faster than cell assays by 1 min. In the genomic DNA assays, DNA
templates were easily accessible to the LAMP reaction components. However, in cell assays,
diffusion of LAMP reaction components inside the cells or their diffusion through cell debris
might have slightly increased the Tt values.

4.3.2 RTf-LAMP assays for quantitative analysis of E. coli and E. faecalis cells

To establish standard curves for quantitative analysis, RTf-LAMP assays targeting uidA
o

gene and gelE gene were performed on commercial PCR instrument (at 63 C) with 10-fold
5

serial dilutions of E. coli and E. faecalis respectively from 10 to 1 cell. Out of 6 positive
controls of single cell assay of E. coli and E. faecalis, only 3 and 4 assays respectively showed
the positive amplification. RTf-LAMP assays of E. coli and E. faecalis showed a consistent
5

increase in the SNR values reaching to a plateau in 15 min for 10 cells and between 20-30 min
for higher dilutions. Tt values showed a consistent increment with respect to decreasing cell
concentrations (Figure 4.2). Negative controls did not show any increase in the SNR throughout
the reaction. For both the assays, standard deviations in the SNR values showed an increase with
respect to decreasing cell concentrations. Additionally, standard deviations in the SNR values of
the E. faecalis cell assays were higher than the E. coli cell assays. It could be attributed to the

133

higher non-specific interactions of Syto-82 with higher amount of cell membranes of E. faecalis.
The standard curves for the RTf-LAMP assays of both the bacterial cells were established by
plotting the Tt values versus log of the number of cell (Figure 4.3). Cell-based RTf-LAMP assays
on commercial PCR instrument were quantitative and sensitive to single cell. Higher standard
deviations in the average Tt values were observed for E. faecalis assay than E. coli assay.
2

Correlation coefficients (R ) of the fitted line were 0.98 and 0.97 for E. coli and E. faecalis
5

assays respectively. Tt values of RTf-LAMP assay from 10 cells to single cell of E. coli were
5

from 8 min to 18.7 min respectively. And Tt values of RTf-LAMP assay from 10 cells to single
cell of E. faecalis were from 6.7 min to 13.3 min respectively. Faster amplification of the E.
faecalis assay was due to rapidly increasing SNR values as compared to E. coli assay (Figure 4.1
and 4.2). Therefore, no preheating was needed for cell-based RTf-LAMP assays i.e., LAMP can
o

be performed by incubating the bacterial cells in the LAMP reagent mixture at 63 C presumably
because enough of the cells are lysed/permeabilized providing the access of genetic contents to
the LAMP solution.

134

Figure 4.2 SNR verses amplification time of the dilution series of RTf-LAMP assays of E. coli
5
(uidA gene) and E. faecalis (gelE gene) cells from 10 cells to single cell on a commercial PCR
instrument. For interpretation of the references to color in this and all other figures, the reader is
referred to the electronic version of this dissertation.

135

4.3.3 Quantification of viable E. faecalis cells by PMA-RTf-LAMP
5

RTf-LAMP assays were performed with 10 E. faecalis (gelE gene) cells treated with
PMA to confirm that the DNA amplification was only from viable cells and not from the dead
cells or extracellular DNA in the solution (Figure 4.4). Differentiation of live and dead cells is a
challenge in DNA-based molecular detection techniques (Girones et al. 2010). Prior treatment of
pathogens with ethidium monoazide (EMA) (Nogva et al. 2003) and PMA (Nocker et al. 2007)
is being applied to test cellular viability by using DNA-based molecular detection techniques.

Figure 4.3 Standard curves of RTf-LAMP assays of E. coli (uidA gene) and E. faecalis (gelE
gene) cells on a commercial PCR instrument

EMA and PMA are photo-reactive dyes with high affinity for DNA. These dyes
intercalate into double stranded DNA and forms covalent linkage upon exposure to blue light,

136

resulting into chemically modified DNA, which cannot be amplified by PCR or LAMP reactions.
These dyes are designed to be cell-impermeable thus targeting only compromised or dead cells
and extracellular DNA, without affecting the viable cells. Comparative investigations of
membrane permeability characteristics of EMA and PMA have shown that EMA could penetrate
the membranes of viable cells leading to cell loss (Nocker et al. 2006; Pan and Breidt 2007).
However, due to higher molecular weight and higher positive charge of PMA than EMA, it
remains membrane impermeant (Yanez et al. 2011). Compared to EMA, PMA has been shown
to have greater success and non-lethal effects in identifying the viability of viruses (Parshionikar
et al. 2010), bacteria (Nocker et al. 2009; Kralik et al. 2010; Chen et al. 2011; Liang et al. 2011),
and protozoa (Sauch et al. 1991; Brescia et al. 2009) coupled with PCR, real-time PCR, and
LAMP assays. Therefore, PMA-LAMP was selected to test the viability of E. faecalis cells. An
optimum concentration 50 µM PMA was selected for this study as a range of concentrations
from 30 to 50 µM was found to be effective in other studies (Pan and Breidt 2007; Vesper et al.
2008). E. faecalis cells treated with PMA were exposure to blue LED for 15 min for the PMADNA complex formation. A longer exposure time (15 min) with low-intensity LED was selected
over a shorter exposure with brighter halogen lamp as heat generated by the halogen lamp has
been reported to kill viable bacterial cells (Wang and Levin 2006; Vesper et al. 2008; DelgadoViscogliosi et al. 2009).

137

5

Figure 4.4 RTf-LAMP assays of 10 E. faecalis non-lysed cells (with and without PMA
treatment) and PMA treated lysed cells on a commercial PCR instrument

RTf-LAMP assays of E. faecalis cells (non-lysed) with and without PMA treatment
showed DNA amplification with the same Tt value of 7 min. SNR value (70.1) at Tt of nonPMA treated cells was higher than the SNR value (50) at Tt of PMA treated cells. Tt and the
corresponding SNR of PMA treated lysed cells were 12 min and 14.5 respectively.

138

Figure 4.5 Colony forming units of non-lysed E. faecalis (with and without PMA treatment). No
growth was observed from heat treated cells

Cell growth was only observed from the non-lysed cells solutions (with and without
PMA treatment) (Figure 4.5). No cell growth was observed from the heat treated cell solution.
RTf-LAMP assays showed a drastic shift in the Tt value of PMA-treated lysed cells than PMAtreated non-lysed cells. A difference in Tt value of 5 min was observed, which corresponded to
4-log fold reduction in the cell concentration due to PMA treatment. No effect of PMA treatment
was observed in RTf-LAMP assays with non lysed cells (with and without PMA treatment).
However, a slight lower SNR value of PMA treated cells than non-PMA treated cells might be
due to death of few viable cells by the heat generated by LED light. Additionally, heat treatment
o

(95 C for 5 min) killed all the viable cells as observed by the lack of any cell growth on TSA
plate (Figure 4.5). These results confirmed that the DNA amplification (Figure 4.1 and 4.2) was
directly from viable bacterial cells rather than dead cells or extracellular DNA in solution.

139

4.3.4 MicroRTf-LAMP assays of single cell of E. coli and E. faecalis

RTf-LAMP assays were performed on polyester microchips and imaged with a CCD
imaging system developed by our group (Ahmad et al. 2011) (Figure 4.6). Polyester material was
compatible with LAMP reaction, as recently reported by our group (Stedtfeld et al. 2011). Also
polyester was comparatively cheaper than cyclic olefin copolymer and no surface treatment of
polyester was needed during/after microfluidic chip fabrication for performing LAMP, thus a
cost effective and easy-to-use option.

Figure 4.6 A polyester chip (2.5 cm × 2 cm) for microRTf-LAMP assays. Each circular well was
1 mm in diameter with a volume of 2 µL. The circular wells were filled with food color

140

Figure 4.7 MicroRTf-LAMP assays of single cell of E. coli (uidA gene) and E. faecalis (gelE
gene). Inset showed the CCD time lapse images of microchip with 10 positive and 10 negative
controls. The error bars represented the standard deviations of the mean from 5 and 8 positive
replicates and 10 negative replicates for E. coli and E. faecalis respectively

141

MicroRTf-LAMP assays with single cell of E. coli (uidA gene) and E. faecalis (gelE
o

gene) were performed at 63 C (Figure 4.7). MicroRTf-LAMP assay of E. coli provided a
maximum average SNR of 438 and an average Tt of 21 min respectively. And microRTf-LAMP
assay of E. faecalis provided a maximum average SNR of 738 and an average Tt of 17.5 min
respectively. Due to dilution effect, only limited number of positive controls showed the DNA
amplification. For example 5 and 8 positive amplifications were achieved out of 10 positive
controls in E. coli and E. faecalis assays respectively. Negative controls did not show any
increase in the SNR throughout the reaction. Higher SNR from E. faecalis assay than E. coli
assay was consistent with the results from commercial PCR instrument. Additionally Tt values
for E. coli and E. faecalis assay on microchip (5 s CCD exposure) was higher by 2.3 min and 4.2
min respectively than the assays on commercial PCR instrument. We had earlier demonstrated
that increasing the CCD exposure time would reduce the Tt value of microRTf-LAMP assays
(Ahmad et al. 2011). However, in this case, reaction wells were covered with mineral oil to avoid
evaporation. Due to the higher refractive index of mineral oil (1.5) than air (1.0), a portion of
emitted photons from reaction wells might have been refracted from the oil-air interface,
reducing the number of emission signals reaching to the CCD and increasing the Tt of cell-based
microRTf-LAMP assays. Also in this case, working distance between the CCD and the
microchip was increased for image focusing, further reducing the incoming emission signals to
the CCD. Tt could be further reduced by increasing the CCD exposure time from 5 s. However,
increasing the CCD exposure time than 5 s would result into poor image contrast, reducing the
image clarity.

142

4.3.5 Quantification of most probable number by digital LAMP
Digital LAMP assays were performed by targeting gadA gene and gelE gene with 10-fold
5

serial dilutions of E. coli and E. faecalis respectively, from 10 cells to 0.01 cell. For each
dilution, 5 replicates were performed to calculate the MPN concentrations of the assays by using
the 10-fold dilution of 5-tube MPN analysis. As LAMP primers for gadA gene of E. coli were
not tested earlier. Therefore, RTf-LAMP assay from single E. coli cell was performed on a
o

commercial PCR instrument at 63 C to confirm the single-cell level sensitivity of the designed
assay. The designed RTf-LAMP assay of E. coli (gadA gene) was sensitive to single cell (Figure
4.8). This RTf-LAMP assay showed slower amplification (Tt: 27 min) than the single cell assay
designed for uidA gene. Therefore, for the MPN-based digital LAMP assays, microchips were
incubated for 40 min before imaging.

143

Figure 4.8 Single cell RTf-LAMP assay of E. coli (gadA gene) on a commercial PCR instrument.
Only two positive controls showed amplification out of six replicates. The error bars represented
the standard deviations of the mean from two replicate experiments

MPN concentrations of single cell LAMP dilutions were quantified as per the statistical
method of result rejection developed by Hurley and Roscoe and Taylor (Taylor 1962; Hurley and
Roscoe 1983). The formula shown below was taken from the analysis of Hurley and Roscoe
(Hurley and Roscoe 1983).

k

m
d ip i vi
 1  exp d xv    d ini vi
i 1
i 1
i
i
Here, k is the levels of dilution, ni is the number of subsamples, pi is the number of positive
subsamples, di is the dilution factor, vi is the volume of each subsample, and x is the estimated
density of the microorganism.

144

The applied assumptions were that bacteria were distributed in the sample without any
clustering and wells containing even a single viable bacterium would produce the
growth/amplification. We have also assumed that the MPN values at the microscale volume
would be consistent with the milliliter volume as used in the developed method. The result
rejection method is based on that if the dilutions provide all the positive results, the calculation
would include only the most diluted samples. If the dilutions provide all the negative results,
calculation would include only the most concentrated samples.
5

Digital LAMP assay of gadA gene of E. coli showed amplification from 10 cells to
single cell (Figure 4.9). However, only 2 out of 5 positive controls wells containing single cell
showed amplification. No amplification was observed with the highest dilutions of 0.1 and 0.01
cell. Based on the 5-well MPN analysis [5 2 0], average MPN concentration of single cell
dilution was 0.49 cell (p = 0.05) per 2 µL of reaction volume. Similarly, Digital LAMP assay of
5

gelE gene of E. faecalis showed amplification from 10 cells to single cell (Figure 4.10).
However, only 3 out of 5 positive controls wells containing single cell showed amplification. No
amplification was observed with the highest dilutions of 0.1 and 0.01 cell. Based on the 5-well
MPN analysis [5 3 0], average MPN of single cell dilution was 0.79 cell (p = 0.05) per 2 µL of
reaction volume. The results implied that the digital LAMP assays had large dynamic range and
single cell-level sensitivity.

145

Figure 4.9 Digital LAMP image for E. coli (gadA gene) dilutions. Cell number per well (left)
and number of positive reaction (right) were shown in the image

Figure 4.10 Digital LAMP image for E. faecalis (gel E gene) dilutions. Cell number per well
(left) and number of positive reaction (right) were shown in the image

146

This approach of digital LAMP would be highly useful for the detection of extremely low
concentration of target cells in higher reaction volume. Digital PCR and digital recombinase
polymerase amplification assays based on DNA amplification have already been reported, which
relies on the readout of high throughput reaction wells (positive/negative) at the end of reaction
(Pohl and Shih 2004; Shen et al. 2011). MPN analysis is generally performed by using culturebased approaches, e.g., broth/agar dilution, which takes 10-12 h for the results (Evans et al.
1981; Rompre et al. 2002; Kinzelman et al. 2003). Digital LAMP could be used for the rapid
bacterial enumeration by gene amplification from cells, thus has the potential to replace the
MPN-based analysis.
4.4 Conclusions

RTf-LAMP assays for rapid and sensitive gene amplification directly from E. coli and E.
faecalis bacterial cells (without any sample processing) were developed. PMA-LAMP was
performed to test the viability of E. faecalis, which also confirmed that DNA amplification was
only from the bacterial cells. RTf-LAMP assays of single bacterial cell were also translated on
microchips. DNA amplification from single cell-level digital LAMP assays was confirmed by
o

MPN analysis. LAMP from bacterial cells at a temperature of 63 C might be due to the cell
lysis or increased cell permeability at this temperature leading to DNA amplification. Direct
DNA amplification from bacterial cells at this moderate temperature would eliminate any sample
processing steps and reduce the complexity of gene analysis instruments. This work would be
helpful towards the simple, sensitive, and rapid integrated gene analysis system with sample-inanswer-out capability.

147

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Indicator of Giardia Cyst Viability. Applied and Environmental Microbiology 57:32433247
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powerful, and smart: Using LAMP for low cost screening of multiple waterborne
pathogen. In K. Sen & N. J. Ashbolt (Eds.), Environmental Microbiology: Current
Technology and Water Applications: Caister Academic Press 103–125
Shen F, Davydova EK, Du WB, Kreutz JE, Piepenburg O, Ismagilov RF (2011) Digital
isothermal quantification of nucleic acids via simultaneous chemical initiation of
recombinase polymerase amplification reactions on slipchip. Analytical Chemistry
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Stedtfeld RD, Tourlousse DM, Seyrig G, Ahmad F, Tiedje JM, Hashsham SA (2011) Gene-Z: A
device for point-of-care genetic testing using a smartphone. Submitted to Lab on a Chip
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Applied Microbiology 25:54-61
Tourlousse DM, Ahmad F, Stedtfeld RD, Seyrig G, Tiedje JM, Hashsham SA (2011) A polymer
microfluidic chip for quantitative detection of multiple water- and foodborne pathogens
using real-time fluorogenic loop-mediated isothermal amplification. Submitted to
Biomedical Microdevices
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viability in air and water samples using quantitative PCR after treatment with propidium
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CHAPTER V
5. RAPID AND SENSITIVE DETECTION OF THE GROWTH OF ESCHERICHIA
COLI AND ANTIBIOTIC SUSCEPTIBILITY TESTING ON MICROCHIP
5.1 Introduction
Gold standard methods for antibiotic susceptibility testing (AST) rely upon culture-based
approaches, e.g., broth/agar dilution and disk diffusion, which are simple and inexpensive (Bauer
et al. 1966; Stalons and Thornsberry 1975). AST for faster growing microorganisms such as
Escherichia coli is performed in 12 -18 h, which increases to 4-6 weeks for slow growing
Mycobacterium species (Limb et al. 1993). These conventional methods are labor intensive and
time consuming to permit same-day diagnosis which delays treatment and increase the risk of
inappropriate use of antibiotics (Jorgensen and Ferraro 2009). Towards the automation and rapid
AST, techniques those measure the bacterial density such as fluorescence (Kim et al. 2010; Joyce
et al. 2011; Sun et al. 2011), optical density (Arora et al. 2009; Chen et al. 2010; Meyer et al.
2011), and micromechanical oscillation (Gfeller et al. 2005) or their activity such as
bioluminescence (Hattori et al. 1998) and radioactive CO2 release (Deblanc et al. 1972) in
microfluidic or microtiter plate formats are being developed. Although the detection times of
these techniques are in hours, they require a number of processing/measurement steps in their
protocols, further delaying the results. Additionally, these techniques are not suitable for pointof-care AST due to the application of expensive detectors such as spectrophotometer,
fluorescence microscope, and cantilever.
Bacterial growth monitoring through fluorescence detection can be used for rapid and
low-cost AST by combining disposable microchips, low-cost and sensitive photon detectors, and
highly fluorogenic membrane staining styryl or Fei Mao (FM) dyes. FM dyes contain two
155

lipophilic tails for partial insertion into the cell membranes and a core containing two aromatic
rings connected with a double bond for generating fluorescence (Gaffield and Betz 2006; Wu et
al. 2009). They provide high fluorescence in the hydrophobic environment i.e., in the presence of
the cellular membranes and remain non-fluorescent otherwise. FM dyes have been extensively
applied for staining prokaryotic (Joyce et al. 2011), eukaryotic (Vida and Emr 1995), and plant
cells (Bolte et al. 2004). A variant of FM dye, FM5-95 can be useful for fluorescence imaging
with simple optical filters and detectors due to the huge difference between the excitation and
emission maxima (Stokes shift of 174 nm). Application of FM5-95 dye for the fluorescence
microscope-based growth monitoring of Mycobacterium species reduced the detection time to 7
h from 4-6 weeks (Joyce et al. 2011). However, to our knowledge there is no report of real-time
monitoring of bacterial growth or AST through fluorescence detection of FM5-95 dye on
microfluidic platform.
o

Here, the real-time growth monitoring of E. coli at 37 C in the presence of a membrane
intercalating FM dye, by using a low-cost microchip and CCD imaging system is reported.
Various concentrations of FM5-95 dye (1 µg/mL to 50 µg/mL) and two different CCD exposure
times (1 s and 10 s) were tested towards the rapid detection of E. coli growth. A combination of
4

20 µg/mL FM5-95 dye and 10 s of CCD exposure time provided the detection of 10 E. coli in
only 1 h. Highest concentration of FM5-95 (50 µg/mL) inhibited the E. coli growth. Sensitivity
4

of the cell growth assay was tested with the dilution series of E. coli from 10 to 1 cell. The
assay was sensitive to single cell with a detection time of 5 h. Antibiotic sensitivity of ampicillin
and tetracycline on this E coli strain was tested. MIC values of ampicillin and tetracycline were
16 µg/mL and 1.6 µg/mL respectively, which were also consistent with the literature values. This

156

method would be useful for high throughput, rapid and real-time monitoring of bacterial growth
for their AST.
5.2 Methods and materials
5.2.1 Microchip fabrication
Microchip with circular wells was fabricated with polyester film. Microchip consisted of
multiple circular wells with 1 mm diameter and 2 µL volume per well. Detailed information
about microchip fabrication process is provided in Chapter IV.
5.2.2 Fluorescence imaging system and signal acquisition
Fluorescence imaging system was built by integrating light source, optical filters, thin
film heater, and a CCD camera. Detailed information about the fluorescence imaging system is
provided in Chapter IV and also in a previous publication (Ahmad et al. 2011). In this modified
system, emission filter (FF01-534/20-25) was replaced with another emission filter (FF01747/33-25, Semrock, Rochester, NY) suitable for FM5-95 dye. Also the heater was maintained at
o

37 C. For real-time monitoring of E. coli growth and AST, microchips were manually imaged
with the CCD at every 5 min till the first 20 min. After 20 min, images were taken at every 20
min for 4-5 h. Light emitting diode was turned on before imaging and then turned off to avoid
any quenching of FM5-95 dye by the constant light illumination. CCD gain of 100, offset of 60,
and exposure times of 1 s and 10 s were used for imaging. Images were saved in flexible image
transport system format. CCD images were opened in Image J (National Institutes of Health,
Bethesda, MD) and a microarray profile plugin (http://www.optinav.com/imagej.html) was

157

applied to obtain fluorescence signals from the images. A microarray grid (circular wells with
diameter of 1 mm) was used to extract the signals from the same pixel region of all the images.
5.2.3 Bacterial cell and genomic DNA
E. coli K12 strain C3000 cells were graciously provided by Professor Joan B. Rose,
Water Quality, Environmental, and Molecular Microbiology Laboratory at Michigan State
University. Frozen stock of E. coli was resuspended in Trypticase Soy Broth (TSB, 211768, BD,
o

Sparks, MD) followed by the incubation at 37 C in a shaker-incubator (New Brunswick
Scientific, Edison, New Jersey) at 150 rpm for 12 h. Broth culture was further inoculated in 25
o

mL of sterile TSB and incubated at 37 C for 8-12 h to achieve mid-log phase. Broth culture was
o

placed overnight in the refrigerator at 4 C. Broth culture was serially diluted in sterile TSB and
colony forming units of E. coli were enumerated by plating 0.1 mL of their dilutions on
Trypticase Soy Agar (TSA, 211043, BD, Sparks, MD) plates. TSA plates were incubated at 37
o

C for 12 h before counting of cell colonies. The concentration of E. coli cells in the broth
8

culture was approximately 1×10 colony forming units/mL. Strict aseptic techniques were
followed throughout the experiments.
5.2.4 Bacterial growth and AST on microchip
Styryl dye, FM5-95, (N-(3-trimethylammoniumpropyl)-4-(6-(4-(diethylamino) phenyl)
hexatrienyl) pyridinium dibromide) was purchased from Molecular Probes (T23360, Invitrogen,
Carlsbad, CA). A stock concentration of 1000 µg/mL of FM5-95 dye was prepared in sterile
o

water and stored at 4 C. The overnight broth culture of E. coli was diluted to approximately

158

3

100-fold in 50 µL fresh TSB to a concentration of 5 ×10 cells/µL. Broth culture of E. coli was
mixed with appropriate amounts of FM5-95 dye solution to prepare cell solutions with dye
concentrations of 50 µg/mL, 20 µg/mL, 10 µg/mL, 5 µg/mL, and 1 µg/mL. A negative control (5
3

×10 cells/µL in TSB without any dye) was also prepared. To establish standard curve for
4

quantitative analysis, 10-fold dilutions of E coli broth culture (10 cells to 1 cell) with 20 µg/mL
FM5-95 dye were prepared. Negative controls containing 20 µg/mL FM5-95 dye in TSB were
also prepared.
Ampicillin (A0166-25G, Sigma Aldrich, St. Louis, MO) and tetracycline (T7660-5G
Sigma Aldrich, St. Louis, MO) were dissolved in sterile water. Stock solutions of ampicillin (50
mg/mL) and tetracycline (10 mg/mL) were filtered through a 0.22-μm nitrocellulose filters
o

o

(Millipore, Billerica, MA) and stored at 4 C and -20 C respectively. E. coli broth cultures with
3

20 µM FM5-95 dye (50 µL; 5 ×10 cells/µL) were mixed with appropriate amounts of ampicillin
to the final concentrations of 32 µg/mL, 16 µg/mL, 8 µg/mL, 4 µg/mL, 2 µg/mL, and 1 µg/mL.
3

And E. coli broth cultures with 20 µM FM5-95 dye (50 µL; 5 ×10 cells/µL) were mixed with
appropriate amounts of tetracycline to the final concentrations of 2.5 µg/mL, 2 µg/mL, 1.5
µg/mL, 1.2 µg/mL, 1 µg/mL, 0.5 µg/mL, and 0.1 µg/mL. Negative controls were free from any
3

antibiotic and only contained 5 ×10 E. coli cells/µL and 20 µM FM5-95 dye in TSB.
4

Approximately 2 µL of broth culture of E. coli (1×10 cells) containing appropriate
amount of dye and or antibiotic was dispensed in each well of the microchip. Each assay was
performed in quadruplicate. Wells were covered with mineral oil (Sigma, St. Louis, MO) during

159

the incubation to avoid the solution evaporation. Afterwards microchip was placed on the heating
o

stage for incubation at 37 C for 4-5 h and real-time CCD-based imaging.

5.2.6 Fluorescence signal analysis
To calculate SNR values of real-time bacterial growth and microAST, raw fluorescence
signals were baseline corrected by subtracting the average fluorescence signals obtained in the
first 15 min and dividing by their standard deviation. Tt was then defined as the time at which
the SNR reached an arbitrary cut-off of 20 or higher. Average of the SNR from the quadruplicate
was plotted with respect to amplification time. Tt values were calculated for individual reactions
of quadruplicate and an average value of Tt is reported here. The error bars in the plots
represented the standard deviations of the mean from four replicate experiments unless otherwise
stated. Data analysis and plotting was done with Microsoft Excel (Microsoft, Redmond, WA).
5.3 Results and discussion
5.3.1 Real-time fluorescence monitoring during E. coli growth

Microchip-based growth of E. coli (1×10

4

cells) in the presence of FM5-95 was

monitored in real-time by measuring the fluorescence signals by a CCD imaging system. To
investigate the effect of dye concentration on the cell growth and the resulting fluorescence
signals, increasing concentrations of dye from 1 µg/mL to 50 µg/mL were applied. FM5-95 dye
has a maximum excitation and emission peak at 560 nm and 734 nm respectively, with a high
Stokes shift of 174 nm. By applying an appropriate emission filter, background light could be
minimized to the detector for highly sensitive detection. Another advantage is the low intrinsic
background from FM5-95 dye. Approximately 40 µM (23 µg/mL) of FM5-95 was found to be
160

optimum, when compared against a variety of styryl dyes towards their brightness and affinity
for liposomal membrane (Wu et al. 2009). We have also included a higher concentration of
FM5-95 dye (50 µg/mL) in this study to investigate its effect on E. coli growth. Two different
CCD exposure times (1 s and 10 s) were applied during imaging to enhance the collection of
photons/fluorescence signals as observed in our microRTf-LAMP assays (Ahmad et al. 2011).
The overall goal was to reduce the detection time of E. coli growth by optimizing the dye
concentration and the CCD exposure time. Fluorescence signals were measured at every 5 min
for the first 20 min (then every 20 min) to have a lower standard deviation of the mean of signal
values in the first 15 min. The measured fluorescence signals were converted into SNR values to
apply the cut-off (SNR≥20) for the calculation of Tt values. Higher SNR values were achieved
due to the lower standard deviation of the mean of signal values in the first 15 min.

161

Figure 5.1 Real-time fluorescence growth curves of E. coli with various concentrations of FM595 dye at 1 s and 10 s of CCD exposure times.

162

E. coli growth curves showed a consistent increase in the average SNR values with
respect the dye concentration and incubation time (Figure 5.1). Fluorescence-based growth
curves exhibited a lag, exponential, and stationary phase similar to real-time optical densitybased measurement of E. coli growth in microchip (Arora et al. 2009). With the CCD exposure
time of 1 s, E. coli with the dye concentrations of 20 µg/mL, 10 µg/mL, and 5 µg/mL showed an
increment in SNR values during the incubation time. The increase in the SNR values with 1
µg/mL dye was relatively lower than higher dye concentrations. The same trend of SNR increase
during the E. coli growth with these dye concentrations was also observed with 10 s of CCD
exposure time. However, no increment in the SNR values was observed with the 50 µg/mL dye,
for both the exposure times. Also the negative controls did not show any increment in the SNR
values during the incubation time. With the CCD exposure time of 1 s, the maximum SNR
values for the 20 µg/mL, 10 µg/mL, and 5 µg/mL dye concentrations were 50.31 ± 2.7, 20.09 ±
1.6, and 2.66 ± 1.9 respectively. And with the CCD exposure time of 1 s, the maximum SNR
values for the 20 µg/mL, 10 µg/mL, and 5 µg/mL dye concentrations were 562.31 ± 68.7, 225.45
± 31.9, and 31.73 ± 6.16 respectively. SNR values with the 10 s of exposure time were
approximately 11-fold higher than the 1 s of exposure time. Our observations were consistent
with other reports demonstrating the improvements in the SNR values with the increasing CCD
exposure times (Yang et al. 2008; Pjescic et al. 2010). Faster increase in the SNR values would
reduce the Tt values leading to faster detection of E. coli growth as microRTf-LAMP assays
(Ahmad et al. 2011).
Lack of increase in the SNR values with 50 µg/mL dye might be due to the E. coli growth
inhibition with such a high dye concentration. In other study, imaging of real-time growth of
Mycobacterium smegmatis on thin agar film was performed with a 10-fold lower FM5-95
163

concentration of 2 µg/mL (Joyce et al. 2011). Also, a maximum concentration of 23 µg/mL of
FM5-95 was applied in the liposomal affinity study (Wu et al. 2009). FM5-95 dye has been
commonly applied for staining cellular membranes of a variety of microorganisms with dye
concentrations ranging from 1 µg/mL to 5 µg/mL (Veach et al. 2004; Adams et al. 2011; Sass et
al. 2011; Willemse et al. 2011). However, cell membrane staining with FM5-95 would not
require the cells to be alive for the endpoint imaging applications. Application of the lower
concentration of FM5-95 dye in the live cell imaging of acid fast gram-positive microorganisms
might be due to the dye-based cell growth inhibition (Joyce et al. 2011). Dye-based growth
inhibition of multiple gram-positive and gram-negative microorganisms with a number of
fluorogenic dyes including, orange II, sudan II, tetrazolium dyes has been reported (Fung and
Miller 1973; Ullrich et al. 1996; Pan et al. 2011). However, no systematic study on the inhibitory
effect of styryl dyes, including FM5-95 is currently available. Growth inhibition might be due to
higher interaction between the FM5-95 and cell membrane E. coli, leading to the inclusion of
dye inside the cells (Bolte et al. 2004). However, this hypothesis was not confirmed in this study.
5.3.2 Effect of FM5-95 dye concentration and CCD exposure time on the Tt value of E. coli
growth on microchip
The effect of FM5-95 dye concentrations (1 µg/mL to 20 µg/mL) and two different CCD
exposure times (1 s and 10 s) on the Tt values was investigated (Figure 5.2). Tt values were
calculated based on the SNR≥20. A higher cut-off of SNR (≥20) was applied, as fluorescence
signals/SNR in these cell assays showed a drastic increase at the starting of exponential phase
(e.g., SNR <2 to ≥20), especially with 10 s of exposure time. With the CCD exposure time of 1 s,
Tt values for 20 µg/mL and 10 µg/mL dye concentrations were 106.67 ± 18.9 min and 133.33 ±
18.9 respectively. Tt for 5 µg/mL dye concentration was not included in the analysis as the SNR

164

with this dye concentration did not increase above the cutoff value. With the CCD exposure time
of 10 s, Tt values for 20 µg/mL, 10 µg/mL, and 5 µg/mL dye concentrations were 60 ± 0.0, 80 ±
11.5, and 100 ± 0.0 respectively. The analysis showed that the Tt values were dependent on both
the dye concentration and the CCD exposure time. Additionally, the fastest detection of E. coli
4

(1×10 cells) growth in 60 min was possibly with a combination of 20 µg/mL FM5-95 dye and
10 s of CCD exposure time.

4

Figure 5.2 Comparison of E. coli (1 ×10 cells) growth detection times (SNR cut-off ≥ 20) with
various concentrations of FM5-95 dye measured with 1 s and 10 s of CCD exposure times.
5.3.3 Quantitative analysis of E. coli dilution series on microchip
4

To establish standard curve for quantitative analysis, real-time growth assays of 10 to 1
cell of E. coli with 20 µg/mL FM5-95 dye were monitored by the CCD imaging system at an

165

exposure time of 10 s. Real-time growth curves showed consistent increase in the SNR values
during the incubation time (Figure 5.3). Fluorescence-based growth curves showed a lag phase,
exponential phase, and stationary phase for all the assays except single cell assay. Higher
incubation time of single cell than 320 min was needed to obtain the complete growth curve. No
increment in the SNR values was observed in the negative controls. Maximum average SNR
values showed a decreasing trend with respect to the decreasing cell number in the assay. The
4

3

2

maximum SNR values for the assays with 10 , 10 , 10 , 10, and 1 cell (s) were 590.3 ± 25.4,
470.4 ± 15.6, and 301.2 ± 15.6, 256.8 ± 16.6, and 165.6 ± 14.6 respectively. Out of 4 positive
controls of single cell assay, only 2 replicate showed the increase in fluorescence signals during
the incubation time of 5 h. That might be due to the dilution effect in single cell assay. Tt values
of the assays were calculated by applying the SNR cutoff value of ≥20. Tt values of assays with
4

3

2

10 , 10 , 10 , 10, and 1 cell(s) were 60 ± 0.0, 120 ± 0.0, 173.3 ± 11.5, 246.7 ± 11.5, and 306.7
±23.1 respectively. The standard curve for growth assay was established by plotting the
logarithmic values of the number of cell versus the Tt values. E. coli growth was quantitative and
2

showed the single cell-level sensitivity. Correlation coefficient (R ) of the fitted line was 0.99.

166

4

Figure 5.3 Real-time fluorescence growth curves of the dilution series of E. coli (1 ×10 cells to
1 cell) with 20 µM FM5-95 dye and the standard curve. For single cell dilution, the error bars
represent the standard deviations of the mean only from 2 replicates

167

Standard growth curve was used for the calculation of the generation time of E. coli
C3000 strain by the formula applied for the growth rate as measured by the increase in the cell
density/optical density with respect to the incubation time . A negative sign was applied as the Tt
values were decreasing with respect to decreasing cell number in the assay. Here, k is the growth
rate, N is the number of cell in the lowest dilution, No is the number of cells in the highest
dilution, and t and to are the Tt values for lowest and highest dilutions respectively.

k = - (ln N - ln No) / (t – to)
4

In our case, the numbers of cell(s) in the lowest and highest dilutions were 10 and 1 and
their respective Tt values 306.7 min and 60 min were respectively. Therefore the growth rate and
generation time of E. coli C3000 strain was 0.037 min

-1

and 27 min respectively. Generation

time of 27 min for this E. coli C3000 strain was earlier reported based on the optical densitybased measurement (Hsu 1968).
5.3.4 Ampicillin and tetracycline susceptibility testing of E. coli in microchip
4

Real-time growth of E. coli (1×10 cells) was monitored in the microchip in the presence
of two different antibiotics, ampicillin (0 to 32 µg/mL) and tetracycline (0 to 2.5 µg/mL), to
examine their inhibitory effect on cell growth. E. coli growth in the presence of 20 µg/mL FM595 dye were monitored by the CCD imaging system with 10 s exposure time under varying
concentrations of ampicillin and tetracycline. Ampicillin and tetracycline were selected in this
study due to their different modes of inhibition on microbial growth (Figure 5.4).
Ampicillin belongs to the penicillin group of β-lactum antibiotics, which acts an inhibitor
for transpeptidase enzyme required for cell wall synthesis, leading to microbial cell lysis
168

(Campolirichards and Brogden 1987). Tetracycline belongs to the family of polyketides also
called protein synthesis inhibitor, which interacts with the 30S ribosome preventing the binding
of aminoacyl-t-RNA to the ribosomal acceptor site, eventually ceasing the microbial growth
(Chopra and Roberts 2001). Both the antibiotics have been found to be effective against a wide
range of gram-positive and gram-negative bacteria, including E. coli (Chopra et al. 1963; Chopra
and Roberts 2001).

a

b

Figure 5.4 Chemical structures of (a) ampicillin and (b) tetracycline

Real-time E. coli growth curves showed a consistent decrease in the average SNR values
with respect the increasing dye concentrations (Figure 5.5). The maximum SNR values of the
growth curves with ampicillin concentrations of 1 µg/mL, 2 µg/mL, 4 µg/mL, 8 µg/mL, 16
µg/mL, and 32 µg/mL were 407.8 ± 25.4, 300.5 ± 15.0, 221.2 ± 15.0, 121.5 ± 20.2, 35.62 ± 6.46,
and 16.6 ± 6.45 respectively. The maximum SNR values of the growth curves with tertacycline
concentrations of 0.1 µg/mL, 0.5 µg/mL, 1 µg/mL, 1.2 µg/mL, 1.5 µg/mL, 2 µg/mL, and 2.5

169

µg/mL were 490.33 ±25.4, 300.5 ± 15.0, 160.4 ± 15.0, 115.55 ± 15.2, 35.58 ± 7.24, 15.56 ± 6.45,
and 15.50 ± 7.24 respectively. The maximum SNR values of the growth curves of negative
controls (with any antibiotic) were 578.30 ± 20.4 and 590.33 ± 25.7 respectively. These SNR
values were consistent with the maximum SNR value of 562.31 ± 68.7 for the E. coli growth
with 20 µg/mL FM5-95 dye and 10 s of CCD exposure time (Figure 5.1). The only difference
between the E. coli growth curves were different slope/rate of growth with the specific type of
antibiotic and their concentrations.

170

a

b

4

Figure 5.5 Real-time fluorescence growth curves of 1 ×10 cells of E. coli (with 20 µg/mL FM595 dye) and increasing concentrations of (a) ampicillin and (b) tetracycline

171

5.3.5 Determination of MIC values of ampicillin and tetracycline for E. coli on microchip
Quantitative evaluation of the inhibitory effects of ampicillin and tetracycline on E. coli
4

(1 ×10 cells) was performed on the microchip. For that, average slopes of the linear portion of
the ampicillin/tetracycline treated growth curves during the exponential phase were calculated.
Data was normalized by dividing the average growth rate with the average slope of the
ampicillin/tetracycline free growth curve (Figure 5.6) (Fesenko et al. 2005; Sun et al. 2011).
Normalized growth rates of E. coli with ampicillin concentrations of 1 µg/mL, 2 µg/mL, 4
µg/mL, 8 µg/mL, 16 µg/mL, and 32 µg/mL were 0.89 ± 0.1, 0.60 ± 0.09, 0.35 ± 0.05, 0.13 ±
0.04, 0.08 ± 0.03, 0.06 ± 0.02 respectively. Normalized growth rates of E. coli with tetracycline
concentrations of 0.1 µg/mL, 0.5 µg/mL, 1 µg/mL, 1.2 µg/mL, 1.5 µg/mL, 2 µg/mL, and 2.5
µg/mL were 0.91 ± 0.1, 0.75 ± 0.09, 0.31 ± 0.09, 0.20 ± 0.09, 0.07 ± 0.03, 0.05 ± 0.02, and 0.04
± 0.02 respectively. The trend showed that the growth rate(s) of E. coli decreased drastically up
to a particular concentration of antibiotic(s) and then became linear (constant) for higher
concentration of antibiotic(s). However, a faster decline in the growth rate of E. coli was
observed with ampicillin as compared to tetracycline. Therefore, ampicillin exerts more acute
inhibition effect on E. coli C3000 than tetracycline.

172

a

b

4

Figure 5.6 Ratio of the growth rates of 1 ×10 cells of E. coli (with 20 µg/mL FM5-95 dye) with
and without (a) ampicillin and (b) tetracycline

173

MIC values of ampicillin and tetracycline were calculated by extrapolating the linear
portion of the growth rate to the antibiotic concentration.

The calculated MIC values of

ampicillin and tetracycline for E. coli C3000 were 16 µg/mL and 1.6 µg/mL respectively.
Inhibitory effects of ampicillin (Chen et al. 2010) and tetracycline (Sun et al. 2011) on other E.
coli strains are being evaluated in microfluidic platforms. The inhibition effect on E. coli growth
in the presence of these two antibiotics and their MIC values, as observed in our case, was
similar to other reports.
5.4 Conclusions
Microchip-based real-time fluorescence monitoring of E. coli growth in the presence of a
novel dye, FM5-95 by using the CCD imaging system was performed. Concentration and CCD
exposure time was optimized for the rapid E. coli growth detection in 1 h. E. coli growth assay
was sensitive to single cell and required a low volume of 2 µL. Effect of E. coli growth
inhibition in the presence of antibiotics were studied. Microchip-based MIC values were found to
be in consistent with the values reported in literature. This real-time, rapid, and sensitive method
of cell growth monitoring could be extended for the AST of microorganisms in high throughput
microchips by applying low-cost CCD imaging system.

174

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175

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CHAPTER VI
6. CONCLUSIONS AND FUTURE PERSPECTIVES
6.1 Conclusions
Culture-based gold standard methods for pathogen/resistance identification are labor
intensive and time consuming resulting into delayed diagnosis of pathogens and increase in
mortality and morbidity. Molecular methods have tremendous potential for addressing the global
threat of infectious diseases due to their rapidity, sensitivity, and accuracy. However, due to their
high cost and complexity, molecular methods are currently used in centralized laboratories.
Therefore, a need exists for a simplified and low-cost platform, on which genetic and cellularbased molecular assays could be performed with similar speed, sensitivity, and accuracy as
commercially available systems. Microfluidic PCR systems are being developed and applied for
gene amplification and detection of a variety of pathogens. However, due to the need for thermal
cycling and the associated high thermal mass, microfluidic PCR systems are slow. Decreasing
the thermal mass and increasing the heating/cooling rates leads to faster gene amplification.
However, that would increase the power consumption and the complexity of the microfluidic
PCR systems. An additional limitation of microfluidic PCR systems is the inhibition of
commonly used Taq polymerase in the presence of cell debris and a variety of other chemicals.
Therefore, completely integrated microfluidic PCR systems (for sample-in-answer-out
capability) require dedicated cell lysis, and nucleic acid purification steps, which increases the
total detection time. Compared to PCR, isothermal nucleic acid amplification methods have not
extensively translated on microfluidic platforms. Translation of loop-mediated isothermal
amplification (LAMP) on microfluidic platform is promising due to i) moderate incubation

180

temperature leading to simplified heating and low power consumption, ii) high amount of
amplification products, which can be detected either by simple detectors, iii) direct genetic
amplification from bacterial cells due to the superior tolerance to substances that typically inhibit
PCR, iv) high specificity, sensitivity up to single copy, and rapid detection in 10-20 min.

Towards the development of a simplified system to perform the rapid detection of
bacterial pathogens and their antibiotic susceptibility, the following key tasks were
accomplished:

1. Developed low-cost polymeric microchips with cyclic olefin copolymer and polyester
and a CCD-based imaging system.
2. Validated microRTf-LAMP assays of 12 virulence genes of 6 diarrheal pathogens on
microchip by using a CCD imaging system. Applied Syto-82 dye and CCD exposure time
control for reducing the Tt of microRTf-LAMP by 10 to 50-fold as compared to
commercial PCR instrument.
3. Developed digital LAMP assays for genetic amplification directly from gram-negative
o

and gram-positive bacterial cells at 63 C without any sample processing.
4. Developed real-time fluorescence-based cell growth assays on microchip by using a
novel cell membrane intercalating dye, FM5-95, for rapid and sensitive detection of E.
coli growth and antibiotic susceptibility testing in 1-2 h.

To our knowledge, the strategies used to reduce the Tt values of microRTf-LAMP assays
and cell growth/microAST method by applying novel fluorogenic dyes and CCD exposure time

181

control, and direct gene amplification from bacterial cells at 63

o

C are novel. These

developments could be easily extended for the rapid testing of pathogens in clinical,
environmental, agricultural, and biodefense applications.
6.2 Future perspectives
In this project, it was shown that gene amplification could be performed directly from
gram-negative and gram-positive bacterial cells in digital format. Also assays for bacterial cell
growth monitoring and AST were developed. Both the assays digital LAMP and AST were
performed separately in this project. However, in future, both the assays will be combined on
single high throughput microfluidic chip for simultaneous detection of pathogens and their
antibiotic susceptibility.

Figure 6.1 Conceptual picture a microfluidic chip for integrated AST and digital LAMP

182

High throughput microfluidic chip will be fabricated by using a low-cost CO2 laser cutter
capable of prototyping high resolution microstructures in polymeric microchips. Microfluidic
chip will have 6 circular patterns containing 18 wells in each circle connected through
capillaries. Also, all the 6 circular patterns will be connected to a central well through capillaries,
o

which will be initially blocked by wax (melting temperature of 65 C). These 6 circular wells
will be used for AST testing (against a variety of antibiotics, as shown below) and then
subsequently for digital LAMP from six individual clinical samples containing diarrheal
pathogens.

183

APPENDICES

184

APPENDICES
Table A1.1 Nucleotide sequences of the designed LAMP primers
Target

Primer

C. parvum
hsp70

F3
B3
FIP

Sequence (5’-3’)
ACAACTCCATCTGTTGTGG

LF
LB

GCATCCCCATTATCAGCC
CAGTGTTTTCTGGGTTTGTGATTGCTTTTTCTGAAGATGGG
CAG
TTACGCAACAAAGAGGCTAATTGGTTATATGGTAAAATTC
CCTGTTCC
TTGCAACTTCACCAACCAATCT
AGAAGGTACGAAGAAGAAGCAATC

F3

ACGCAGAAGGCAGTCAAGA

B3
FIP

ACCACACTTCAATGTCGCAG
GGAAGCAGCACTAGTTTGGCCAACTGAAGCTTCTGGTAGC
CA
TCCAGCTCAAAGTGAAGGCGCGGGGTACCTTCTCCGAACC
TCACTACCTTCCTCTTCAGAACC
TGCGGCACTTCATTTGTAATGT

BIP

C. parvum
gp60

BIP
LF
LB
S. enterica
invA

F3

LF
LB

GAAGCGTACTGGAAAGGGAA
TCAACAATGCGGGGATCTG
ATGATGCCGGCAATAGCGTCACAGCCAGCTTTACGGTTCC
T
GATGACCCGCCATGGTATGGATACCATCACCAATGGTCAG
C
TGATAAACTTCATCGCACCGTC
TTRTCCTCCGCTCTGTCTAC

F3

GCCATTCCACATCGAAGAGGT

B3
FIP
BIP
LF
LB

ATGAGAACATCAATGGTATGGC
GGCGTGAGAGATCCACCTGGAATGCGCCGTAATAGCGGTC
CACCATTATGGAAACGCTTATCCGCCGGATACAGCTGAAG
CATC
CAGGTGATCAACATCCCGCC
CGGTAAAGTGGTCAGCAAAGAT

F3

CCCCACCTTTAACTAGAACAA

B3

GCCAAATCCTTTGCTATCGA

B3
FIP
BIP

S. enterica
phoB

C. jejuni
cdtA

185

Table A1.1
(cont’d)
FIP
BIP

C. jejuni
0414

LF
LB
F3
B3
FIP

CGCAGGGCGATTTTCAAAAATTAAACAATAATGCAGYAAA
TGGGAT
TTAACCATTTTAGGCCCTAGCGGCAAATCCAATTTCCTTGT
GCT
GCTTCGTCTTTAAAGYGAGGATTG
AGCAGCTTTAACGGTTTGGG
GCAAGACAATATTATTGATCGC

LF
LB

CTTTCACAGGCTGCACTT
ACAGCACCGCCACCTATAGTAGAAGCTTTTTTAAACTAGG
GC
AGGCAGCAGAACTTACGCATTGAGTTTGAAAAAACATTCT
ACCTCT
CTAGCTGCTACTACAGAACCAC
CATCAAGCTTCACAAGGAAA

F3

AGCTCTAACAATGTACAGCT

B3
FIP
BIP
LF
LB

AGTTGCAGTTCCTGAAACA
GTCTTATCCGCCGTAAAGTCCGCCGTTCTGTCGAATGGTC
CTAAAGCGGATAACGCCGATACCCAGGGACATTAGCCTGA
G
CCCAACCTGGTCGACAACTT
ATTACTTATACCGCGACGGTGAA

F3

GAGATATCGACCCCTCTTG

B3
FIP
BIP
LF
LB

AATCTGAAAAACGGTAGAAAGT
TCCACAGCAAAATAACTGCCCAACATATATCTCAGGGGAC
CA
GATGTCTATCAGGCGCGTTTTGCCGTATTAACGAACCCGG
TGTGGTTAATAACAGACACCGATG
ACCATCTTCGTCTGATTATTGAGC

F3

TCGGGCAGATTCTAGACC

B3
FIP

GTGGGCACTTCTCAAACT
TTGAGTACCTCGGTCAAAGTACTTCCTGATGAAATAAAGC
AGTCA
TCAACCTTTATGATCATGCAAGAGGGGAAACATATCCATC
ATCGTG

BIP

E. coli
O157:H7
eaeA

E. coli
O157:H7
stx2

V.
cholerae
ctxA

BIP

186

Table A1.1
(cont’d)
CCTCTTGGCATAAGACCACC
AACTCAGACGGGATTTGTTAGG
CGAGTGGAAACGGTTGAAGA

B3
FIP

LF
LB

AGGGGAAGTAAGACCGCTAT
GCACACTGCTTGAYTCTGCGTACGAAAGCGAAGCTGCTCA
T
AGCCACTGTAGTGAACACACCGTCGATTCCCCAAGTTTGG
AG
ACAGATTCTGGCTGAGAGATGTC
CAGCCAGCCAATGTTGTGAC

F3

GTCAAGCAACCTGATCT

B3
FIP

LF
LB

TGCCGCAATCAAAATCCT
CCA
TCATGGAATTGTTGACAAATCCTAATCCTCAACGACAGTCT
G
GCCAGGTCAACCAGGAATAAAACGCTTCATATAATAAAG
CGACGT
GAGGACAATGGCCCACTCA
CGTTGACCTTTGCCAATCTGA

F3

TCTTTACGTTGATGAACTAGC

B3
FIP
BIP
LF
LB

AAGTTGAAAGTATTAGGGTCTAC
AGAGCAAACAGATTAGGGGTGCCATCACGTGTGAAGCCAA
ATAGCGGTGGTGAAGCGATTGTAAATAGGCGGCGCAAT
TGTCTATTCTTTCTTTTGCCCAGAC
CACGGACATGCCTTTGTTCC

F3

CKGTAGAAACCCCAACCCG

B3
FIP
BIP
LF
LB

V.
cholerae
toxR

LF
LB
F3

AWACGCAGCACGATACGC
TAACGCGCTTTCCCACCAACGGCCTGTGGGCATTCAGTC
TAACGATCAGTTCGCCGATGCACTGCCCAACCTTTCGGTAT
TCCACAGTTTTCGCGATCCA
ACGTCTGGTATCAGCGCGAAGT

F3

GCTGTTCTGCTGGGCAATA

BIP

L.
pneumophi
la dotA

BIP

L.
pneumophi
la lepB

E. coli
uidA

E. coli
gadA

187

Table A1.1
(cont’d)
B3
FIP
BIP
LF
LB
E. faecalis
gelE

CGGGATACACCCTGTACGA
GTCGCGGCTTCGAAATGGACTTTTCGGGTGATCGCTGAGA
GCATCACCACGATGTCGGTGGTCTCTGAACGTCTGCGTCT
AGACTACAAAGCCTCCCTGAAATA
CTTCACCGCCGAGAGTGA

F3

GTTGATGAACAACATCCAGA

B3
FIP

CATTCAACGCACCTGATTG
CGTTTTTCCTGTTGGTGTACTTGTTTGCTTATGACAATGCTT
TTTGG
GCTTCCTCTTTAGATGTAGTTGGTCTAAATATTCTAAACCG
GCAGT
CCATAACGCATTGCTTTTCCATC
CATGAAATGACACATGGTGTAACGG

BIP
LF
LB

188