DEVELOPMENT OF A MULTI-MODE OPTICAL IMAGING SYSTEM FOR
PRECLINICAL APPLICATIONS IN VIVO
by
Jae Youn Hwang
A Dissertation Presented to the
FACULTY OF THE GRADUATE SCHOOL
UNIVERSITY OF SOUTHERN CALIFORNIA
In Partial Fulfillment of the
Requirements for the Degree
DOCTOR OF PHILOSOPHY
(BIOMEDICAL ENGINEERING)
August 2009
Copyright 2009 Jae Youn Hwang
ii
DEDICATION
I dedicate this thesis to my lovely wife, Hayong Kim, my parents, Moon Sung Hwang
and Myung Lim Jeong, my sisters, Jeeyoung Hwang and Joohyun Hwang, and my
parents-in-law, Inkil Kim and Eunja Lee.
Without their patience, understanding, support, and most of all love in many difficulties,
the completion of this work would not have been possible.
iii
ACKNOWLEDGEMENTS
During these graduate years at the University of Southern California, I have been
tremendously blessed to be accompanied and supported by many people directly or
indirectly. The completion of this dissertation would not have been possible without them.
It is a great pleasure that I have a brilliant opportunity to thank all of them.
First, I would like to gratefully and sincerely thank to my advisor, Dr. Daniel L. Farkas
for his guidance, understanding, and patience during my graduate studies. His mentorship
was paramount in providing a well-rounded experience consistent my long-term career
goals. He encouraged me to not only grow as a scientist but also as an independent
thinker. Also, he has kept consistent interest in the progress of my thesis and was
available when I needed his expertise and advice.
Special thanks also go to Drs. Lali Medina-Kauwe, Harry B. Gray and Zeev Gross for
much of their support for my research. Drs. Kevin Burton and Erik Lindsley provided
helpful and perceptive comments. In addition, I would like to thank all my colleagues at
the Minimally Invasive Surgical Technologies Institute in Cedars-Sinai Medical Center
for their kindness and friendship. Among them, V. Krishnan Ramanujan deserves special
thanks, as does Dr. Sebastian Wachsmann-Hogiu, now at University of California at
Davis.
Finally, I would like to thank my lovely family, my wife, my parents, and parents-in-law.
Thank you for all your understanding, patience, and support, Love you all.
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TABLE OF CONTENTS
DEDICATION.................................................................................................................... ii
ACKNOWLEDGEMENTS............................................................................................... iii
LIST OF TABLES............................................................................................................. vi
LIST OF FIGURES .......................................................................................................... vii
ABBREVIATIONS ........................................................................................................... xi
ABSTRACT..................................................................................................................... xiii
CHAPTER 1: INTRODUCTION....................................................................................... 1
1.1. Small animal imaging ...........................................................................................1
1.2. Imaging modalities for small animals...................................................................2
1.3. Necessity of a multimode optical imaging system ...............................................3
1.4. Limitations of current commercial systems ..........................................................5
1.5. Significance and Hypothesis.................................................................................6
1.6. Dissertation Organization .....................................................................................7
CHAPTER 2: BACKGROUND....................................................................................... 10
2.1. Optical properties of fluorophores ......................................................................10
2.2. Optical imaging techniques for small animal imaging .......................................15
2.3. Effects of the tissue window on both excitation and emission photons..............23
2.4. Chemotherapy molecules for cancer treatments .................................................24
CHAPTER 3: METHODS AND MATERIALS .............................................................. 27
3.1. Multimode optical imaging system.....................................................................27
3.2. Imaging modes in a multimode optical imaging system ....................................34
3.3. Wide-field two-photon excitation for multimode optical imaging in vivo .........42
3.4. Experimental setup and materials for investigation of optical characteristics
of S2Ga ...............................................................................................................46
3.5. Fluorescence lifetime imaging of HerGa and S2Ga on human cancer cells.......49
3.6. Measurement of mitochondria membrane potential variations of the breast
cancer cells induced by HerGa ...........................................................................50
3.7. Spectral imaging with a ratiometric analysis method for dynamic monitoring
of HerGa accumulation and clearance ................................................................51
v
CHAPTER 4: RESULTS.................................................................................................. 54
4.1. Single modality experiments...............................................................................54
4.2. The feasibility of multimode optical imaging in vivo for simultaneous
complementary/different information.................................................................75
4.3. Multimode optical imaging in chemotherapy assessment and cancer
detection..............................................................................................................77
CHAPTER 5: CONCLUSIONS AND DISCUSSION..................................................... 99
BIBLIOGRAPHY........................................................................................................... 110
APPENDIX A: System control programs....................................................................... 122
A.1. Main system control program ..........................................................................122
A.2. Scanning/wide-field two-photon control program...........................................124
A.3. Joystick/mFLIM program ................................................................................125
A.4. 3D fluorescence imaging program...................................................................126
A.4. Source code ......................................................................................................126
APPENDIX B: Fluorescence lifetime imaging and analysis.......................................... 127
B.1. First-order exponential fitting method for fluorescence lifetimes....................127
B.2. Source code ......................................................................................................127
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LIST OF TABLES
Table 1-1. Selected modalities for small animal imaging....................................................3
Table 2-1. Peak excitation and emission wavelength of fluorophores. .............................12
Table 8-1. Application and capabilities of each imaging mode in the multimode
optical imaging system. ....................................................................................99
Table 8-2. Multimode optical imaging in assessment of nanoconstructs and HerGa
chemotherapy..................................................................................................100
vii
LIST OF FIGURES
Figure 1-1. Functionality of multimode optical imaging system.........................................6
Figure 2-1. Jablonski diagram of excitation and emission, and the corresponding
excitation and emission spectra.......................................................................12
Figure 2-2. Fluorescence intensity image ..........................................................................15
Figure 2-3. Spectral imaging..............................................................................................17
Figure 2-4. Fluorescence lifetime imaging ........................................................................19
Figure 2-5. Jablonski diagram and geometry of two-photon excitation. ...........................21
Figure 2-6. Scanning method through a fiber bundle and the applications of the intra-vital
confocal mode (from Mauna Kea Technologies, Paris, France).............22
Figure 2-7. Chemical reaction process of luciferase enzymes and enzyme-specific
substrates (luciferin)........................................................................................23
Figure 2-8. Optical window for tissue imaging. ................................................................24
Figure 2-9. Schematic of Nanoconstructs (Polycefin) (Ljubimova et al., 2008)...............24
Figure 2-10. Chemical structure of sulfonated gallium corroles (from Harry B. Gray)....26
Figure 3-1. Multimode optical imaging system.................................................................28
Figure 3-2. Laser delivery system......................................................................................30
Figure 3-3. Relay lens system for filter installation...........................................................32
Figure 3-4. System control program ..................................................................................33
Figure 3-5. Overall system control scheme. ......................................................................33
Figure 3-6. Fluorescence imaging mode............................................................................34
Figure 3-7. Spectral imaging mode....................................................................................35
Figure 3-8. Mosaic fluorescence lifetime imaging using fs pulsed laser...........................38
viii
Figure 3-9. Intra-vital confocal imaging mode ..................................................................40
Figure 3-10. Bioluminescence imaging mode ...................................................................40
Figure 3-11. Scanning/wide-field two-photon imaging modes .........................................42
Figure 3-12. Schematics of the experimental set-up for wide-field two-photon
excitation .......................................................................................................45
Figure 3-13. Experimental setup for corrole fluorescence lifetime imaging .....................47
Figure 3-14. Composite spectra with varying ratios of spectra of autofluorescence and
HerGa fluorescence. ......................................................................................52
Figure 4-1. Fluorescence intensity images of a mouse injected with drug molecules
tagged by rhodamine B and Alexafluor 680 ...................................................55
Figure 4-2. Contrast comparison between images recorded with a commercial system
and our multimode optical system ..................................................................56
Figure 4-3. Spectral imaging of a nude mouse with injection of fluorescein into
implanted tumors.............................................................................................57
Figure 4-4. Large area image obtained by mosaic fluorescence lifetime imaging ............59
Figure 4-5. Fluorescence lifetime image of a nude mouse with an implanted breast
tumor ...............................................................................................................60
Figure 4-6. Intra-vital confocal imaging of the rat spine ...................................................61
Figure 4-7. Bioluminescence image from engineered mice (normalized by highest
value)...............................................................................................................61
Figure 4-8. Two-photon excited fluorescence image of a mouse liver stained with
fluorescein (Nikon 20x, NA: 0.50) .................................................................62
Figure 4-9. Optical characteristics of wide-field two-photon excited fluorescence
imaging............................................................................................................65
Figure 4-10. Comparisons of depth dependence in one-photon, scanning two-photon,
and wide-field two-photon excited fluorescence imaging of a tissue
phantom.........................................................................................................66
Figure 4-11 Wide-field two-photon excited fluorescence images of a mouse intestine....67
ix
Figure 4-12. One-photon and wide-field two-photon excited fluorescence imaging of
an eye specimen of an alzheimer’s mouse ....................................................68
Figure 4-13. Wide-field two-photon excited fluorescence imaging and multimode
optical imaging with wide-field two-photon excitation of a nude mouse
with implanted tumors after subcutaneous injection of corroles and
miscrosphere solution....................................................................................70
Figure 4-14. Multimode optical images of a mouse with corroles injected into an
implanted tumor region and the middle of back ...........................................76
Figure 4-15. Fluorescence intensity changes within tumor region and background .........79
Figure 4-16. Clearance examination of the drug nanoconstruct using spectral imaging...80
Figure 4-17. Quantitative images of organs.......................................................................81
Figure 4-18. Concentration dependence of fluorescence lifetime of 50 μM S2Ga in pH
7.5 PBS (room temperature)..........................................................................83
Figure 4-19. pH dependence of fluorescence lifetime of S2Ga at different
temperatures ..................................................................................................83
Figure 4-20. Fluorescence lifetime changes of S2Ga due to temperature changes ...........84
Figure 4-21. Fluorescence of corrole (50uM) versus excitation wavelength (power :
300mW, emission : 620nm) ..........................................................................85
Figure 4-22. Fluorescence lifetime changes of 50μM S2Ga (at pH 7.5 in PBS) on
breast cancer cells (MDA-MB-435) at 36.5ºC..............................................86
Figure 4-23. Fluorescence lifetime changes of HerGa due to glioma cancer cell
endocytosis ....................................................................................................87
Figure 4-24. Measurement of mitochondrial membrane potential of the breast cancer
cells using TMRM fluorescence imaging .....................................................88
Figure 4-25. HerGa distribution in nude mice ...................................................................90
Figure 4-26. Quantitative examination of HerGa accumulation kinetics and clearance
from a mouse at different time points using spectral imaging with
ratiometric and standard spectral classification ............................................92
Figure 4-27. Fluorescence lifetime image and a lifetime histogram of the mouse at 4
day .................................................................................................................94
x
Figure 4-28. Specific organs and tumor images obtained using multi-mode optical
imaging..........................................................................................................95
Figure 4-29. Multimode optical imaging for cancer detection and delineation of
HerGa ............................................................................................................97
Figure A-1. Port number panel ........................................................................................122
Figure A-2. Main system control panel ...........................................................................123
Figure A-3. Server connection panel ...............................................................................124
Figure A-4. Scanning two-photon excited fluorescence imaging panel ..........................125
Figure A-5. Joystick/FLIM panel ....................................................................................125
Figure A-6. 3D fluorescence imaging panel ....................................................................126
xi
ABBREVIATIONS
AMP : Adenosine monosphate.
AON : Antisense oligonucleotides.
ATP : Adenosine triphosphate.
BBO: Barium-borate.
CCD : Charge-coupled device.
CT : Computed tomography.
DIC : Differential interference contrast.
FLIM : Fluorescence lifetime imaging.
FRET : Fluorescence resonance energy transfer
fs : Femotosecond.
H&E : Hematoxylin and eosin (predominant stain in histopathology)
HerPBK10 : Breast cancer-targeted carrier protein.
HerGa : Targeted sulfonated gallium corrole, HerPBK10 + S2Ga.
LCTF : Liquid crystal tunable filter.
LED : Light-emitting diode.
MRI : Magnetic resonance imaging.
Nanoconstruct : Novel platform for tumor treatment (Polecefin).
PBS : Phosphate buffered saline.
PEG : Polyethylene glycol.
PET : Positron emission tomography.
PMT : Photomultiplier tube.
xii
PMLA : Poly β-L-malic acid.
ps : Picosecond.
RSSE : Root sum of square error method.
SPECT : Single photon emission computed tomography.
S2Ga : Sulfonated gallium corrole.
TMRM : Tetramethyl rhodamine methyl ester.
TMRE : Tetramethyl rhodamine ethyl ester.
US : Ultrasound.
xiii
ABSTRACT
This thesis reports advanced optical imaging technology in a stand-alone system that
enables functional mesoscopic imaging (whole-body or endoscopic imaging with
microscopic resolution) of small animals in vivo, and provides for quantitative, dynamic,
and functional monitoring of chemo- and nanoconstruct therapy. Many currently
available imaging approaches have been applied to preclinical studies of cancer, stem
cells, and pharmaceutical outcomes. Moreover, useful in vivo imaging may require
several, preferably combined, and advanced imaging modalities to examine different but
complementary characteristics of molecules, cells, or tissues. Although commercial
systems perform well for standard imaging of small animals, they have limitations
stemming from being single-modality instruments. Thus, a new multimode optical
imaging system that is designed to be application-optimizable, with higher sensitivity and
specificity has been developed here in order to overcome these limitations. The
instrument combines various in vivo imaging modes, including fluorescence intensity,
spectral, lifetime, intra-vital confocal, two-photon excited fluorescence, and
bioluminescence imaging. Also this system is a unique and comprehensive imaging
platform for analyzing kinetic, quantitative, environmental, and other highly-relevant
information with macro- to micro-scopic resolution. This system can be optimized for
various applications, and the combination of multiple imaging modes for increased
contrast and complementary/synergetic information in chemotherapy assessment and
cancer detection is emphasized here.
1
CHAPTER 1
INTRODUCTION
1.1. Small animal imaging
Biomedical research has translational aims, consisting in moving laboratory-derived
methods, experiences and understanding into the clinic. This has shifted emphasis from
cells alone to higher levels of biological organization, and enhanced the use of animal
models with human disease phenotypes such as cancer and Alzheimer’s (Spires and
Hyman, 2005). Such models play a very significant role in preclinical studies since these
allow the detection of primary and metastatic tumors, as well as monitoring the effects of
repeated pharmacological intervention (Gleysteen et al., 2008; McElroy et al., 2008;
Takeda et al., 2008). In the past, it was necessary to sacrifice an animal in order to detect
the tumors or monitor the drug molecules inside the tissues. The excised tissues are not
functional, and thus the normal physiology such as blood flow, cell-cell interactions, and
metabolic activity are suspended, altered, or degraded. This makes it impossible to
continuously monitor biochemical, genetic, and pharmacological processes in vivo and in
the same animal, and in a repeated fashion. Thus, better methods to track disease
development and to monitor tissues and molecules within living animals are needed. It is
for this purpose that in vivo small animal imaging methods are being developed. The
small animal imaging allows us to measure and analyze anatomical, functional, and
molecular processes noninvasively and repeatedly in entire animals. Thus, these new
2
methods go beyond the study of thin tissue sections or cell cultures to technologies that
can operate in vivo with high sensitivity and specificity (Contag, 2007; Graves et al.,
2003; Rice et al., 2001; Weissleder, 2002).
1.2. Imaging modalities for small animals
Currently, many non-invasive technologies such as ultrasound (US), X-ray, computed
tomography (CT), magnetic resonance imaging (MRI), positron emission tomography
(PET), single photon emission computed tomography (SPECT), and optical imaging are
available for small animal imaging. These were originally developed for clinical
applications, but have been redesigned for small animal imaging. US, CT and MRI are
capable of resolving the anatomy and physiology through the energy-tissue interaction.
On the other hand, PET and SPECT require reporter probes or contrast agents in order to
monitor metabolism of organisms (Contag, 2007; Diamandis, 1991; Mahmood and
Weissleder, 2003; Rice et al., 2001; Weissleder, 2002). In particular, various optical
imaging technologies which range from 2D imaging to 3D tomography of internal organs
and tissues have been developed for small animal imaging (Davis et al., 2008; Deniset-
Besseau et al., 2007; Ryu et al., 2008; Soubret and Ntziachristos, 2006; Zacharakis et al.,
2005). These methods have several advantages over the other modalities previously
mentioned (Frostig, 2002). They can offer simultaneous monitoring of multiple targets or
molecular pathways, and are relatively cheaper and less harmful than the non-optical
modalities (Contag, 2007; Diamandis, 1991; Rice et al., 2001; Weissleder, 2002).
3
Technique Resolution Primary use
MRI ~100 μm
US ~50 μm
X-ray ~50 μm
Anatomy/ Physiology
PET ~2000 μm
SPECT ~2000 μm
Metabolism/ Molecules
Optical Imaging ~1 μm Molecules
Table 1-1. Selected modalities for small animal imaging.
1.3. Necessity of a multimode optical imaging system
Preclinical research including cancer, chemotherapy, stem cell, and toxicological study
may require single or multiple optical imaging techniques, depending on their application
(Brinkmann et al., 2008; Kaijzel et al., 2009; Medarova et al., 2006). The optical imaging
techniques, including single-photon, single wavelength fluorescence intensity, spectral,
fluorescence lifetime, intra-vital confocal, two-photon, and bioluminescence imaging,
have the inherent capability to provide different information; fluorescence intensity
imaging can provide either kinetic or dynamic information through monitoring the
changes of fluorescence intensity of molecules stained with fluorophores (Almutairi et al.,
2008; Levitt et al., 2006); spectral imaging allow us to obtain better quantitative
information than the fluorescence intensity imaging (Li et al., 2008; Stamatas and Kollias,
2007; Thaler and Vogel, 2006; Wells et al., 2007; Burton et al., 2009); fluorescence
lifetime imaging can provide functional and environmental information around
fluorophores since the fluorescence lifetime of fluorophores is extremely sensitive to
4
their surroundings (Bloch et al., 2005; Elson et al., 2004; Jo et al., 2005; Nothdurft et al.,
2009; Ramanujan et al., 2008); the two-photon excited fluorescence and intravital
confocal imaging can provide high resolution information of cellular and molecular
events from intact tissues or small animals in vivo (Contag, 2007; Tirlapur and Konig,
2001).
In particular, in chemotherapy research, a variety of assessments of the chemotherapy
drug molecules are required such as drug targeting capability (fluorescence intensity
imaging) (Agadjanian et al., 2009; Ljubimova et al., 2008), in vivo clearance of the drugs
(spectral imaging), and the effects on tissues and organs (spectral, lifetime, and two-photon
excited fluorescence imaging) (Han et al., 1995). Here, it is necessary to utilize
appropriate optical imaging techniques for accurate examinations. Furthermore, a
combination of multiple optical imaging techniques can provide better assessment of
drugs (Bullen, 2008; Cai et al., 2008; Ljubimova et al., 2008; Medarova et al., 2009;
Wang et al., 2007; Zhang et al., 2008).
Finally, a multimode optical imaging platform, combining novel techniques for clinical
applications may enhance cancer detection, diagnosis, decision-making, and treatment
outcomes during surgical procedures (Balas, 2001; de Ryk et al., 2008; Feifel and
Hildebrandt, 1992).
5
1.4. Limitations of current commercial systems
A number of companies have developed optical imaging systems for small animal
imaging. Although the commercial systems perform well for the tasks they have been
designed for, they have limitations that stem from being single modality instruments
(fluorescence or luminescence intensity, or intravital confocal imaging). However, a
variety of biomedical research may require the combination of several advanced imaging
modalities to examine different but complementary characteristics of molecules, cells,
and tissues. Currently, there is no such multimodal optical imaging system commercially
available. Although it is possible to carry out small animal imaging with each commercial
instrument separately, the different times, locations, and animal positioning (in addition
to costs and logistics) may hinder efficacy.
These limitations, and the need we see for more versatile and advanced in vivo imaging,
prompted our development of a multimode optical imaging system, based on the
multimode concept originally introduced for microscopy in the early nineties (Farkas et
al., 1993). In this dissertation, a new multimode optical imaging system designed to be
application-optimizable, with higher sensitivity and specificity, is developed, and the
application to chemotherapy research is described. This instrument combines various in
vivo imaging modes, including one or two-photon excited fluorescence, spectral,
fluorescence lifetime, intra-vital confocal, and bioluminescence imaging.
6
1.5. Significance and Hypothesis
This multimode optical imaging system employs a synergetic combination of multiple
imaging modes that exploit the combined advantages of fluorescence, spectral, high
sensitivity lifetime, two-photon excited fluorescence, and intra-vital confocal detection of
small animals in vivo. Figure 1-1 shows the functionality and capability of each imaging
technique. A successful completion of this system implies the following outcomes: (i)
designing an innovative imaging platform, (ii) establishing multimode imaging as a new
pre-clinical research tool, (iii) monitoring chemotherapy drug dynamics and targeting
efficacy, and (iv) building a stand-alone system for various applications and a common
work flow of small animal studies with all optical, hardware, and software components
needed.
Figure 1-1. Functionality of multimode optical imaging system.
Most importantly, in chemotherapy research, the combination of the imaging modes can
be established as a powerful mesoscopic imaging method for dynamic, quantitative, and
functional monitoring of drug molecules and their discrimination with macro- and micro-
Kinetic/dynamic Functional
Endoscopic high resolution
Multimode optical
imaging system
Spectral
Intra-vital
confocal
Bioluminescence
Fluorescence
lifetime (FLIM)
Two-photon excited
fluorescence
High sensitivity High resolution
and penetration
Quantitative
Single-wavelength,
single-photon
intensity
7
scopic resolution. In addition, these capabilities will enhance the imaging discrimination
ability of normal and diseased tissues in vivo. Furthermore, when the multimode optical
imaging is adapted to endoscope techniques, it may provide better detection, diagnosis,
navigation, decision-making, and treatment during surgical procedures, more closely
linking – spatially and temporally – diagnosis and intervention.
1.6. Dissertation Organization
In Chapter 2, background for a multimode optical imaging system will be introduced. A
variety optical imaging techniques is described in detail. Also, nanoconstucts and
targeted sulfonated gallium corroles among chemotherapy molecules will be described
because we use these molecules in this thesis.
In Chapter 3, methods and materials are described. The development of the multimode
optical imaging system is introduced. A schematic of the multimode optical imaging
system is demonstrated, as well as functionalities (laser delivery, fluorescence detection
and possible imaging modes) and operation strategy (operation procedures and program)
of the system is described in detail. Also, each imaging mode in the multimode optical
imaging system is described.
In Chapter 4, results are described. Here, the evaluation of each imaging mode, including
fluorescence intensity, spectral, lifetime, intra-vital confocal, bioluminescence, and
scanning/wide-field two-photon excited fluorescence imaging, is performed with in vivo
and in vitro specimens, and the feasibility of the multimode optical imaging is examined
8
in chemotherapy study. Additionally, the wide-field two-photon excitation method for
live small animal and multimode optical imaging in vivo using non-scanning imaging
devices is demonstrated because of it’s novelty and the usefulness in the multimode
optical imaging system. As one of applications of multimodality, multimode optical
imaging for the assessment of nanoconstucts in breast cancer treatments is introduced.
Here, the targeting capability and effects of the nanoconstruct are evaluated using
fluorescence and spectral imaging mode with spectral unmixing. As another application
of our system, the power of combining advanced optical imaging technologies
(fluorescence intensity, spectral, lifetime, and two-photon excited fluorescence imaging)
is brought to cancer detection and treatment assessment of targeted gallium corrole
(HerGa). Here, basic optical characteristics (fluorescence lifetime and two-photon
excitation wavelength) of the sulfonated gallium corrole (S2Ga) are investigated. Also, in
order to investigate cell mechanism of the corroles, the fluorescence lifetime changes of
HerGa and S2Ga due to the endocytosis of cancer cells are monitored in real-time as well
as mitochondrial membrane potential variations of the cancer cells, which indicate a
degree of health status of cells, induced by the corroles are examined. In addition,
multimode optical imaging is performed for the assessment of HerGa in breast cancer
chemotherapy. Furthermore, the feasibility of cancer detection and delineation by using
multimode optical imaging of HerGa is examined. Finally, spectral imaging with
ratiometric analysis method is introduced for monitoring dynamics and kinetics of
chemotherapy drug in vivo. Particularly, the examination of HerGa accumulation kinetics
in nude mice following intravenous injection of HerGa is described using the spectral
imaging with ratiometric analysis method.
9
In Chapter 5, concluding comments and discussion of ideas for future research are
presented.
10
CHAPTER 2
BACKGROUND
2.1. Optical properties of fluorophores
Fluorophores have been used as key molecules in optical imaging. The fluorophores can
emit light with their inherent wavelength. If fluorophores are exposed to light, the light is
absorbed by the fluorophores. And then, an electron in the ground state of the fluorophore
is promoted to an electronically excited state. Return to the ground state is accompanied
by emission of light. This process is referred to as fluorescence.
2.1.1. Principle of fluorescence
Fluorescence is light emitted from a molecule. The release of a photon into the
environment is a response of a molecule to an excited state. There are four common
methods by which the excitation energy is delivered to the molecule: (1) Impinging
photons, which are a major source for excitation of a molecule. If photons impinge a
molecule such as a fluorophore, it can cause the excitation of the molecule; (2) Change
shift (chemical energy). Chemical reaction of molecules can produce chemical energy
[adenosine triphosphate (ATP)]. If the chemical energy is absorbed by a molecule, the
molecule can be excited; (3) Thermal photons, which can be absorbed by molecules for
excitation. If the thermal photons have enough energy to excite molecules, the molecules
11
are excited; (4) Resonance effects. If two different molecules are located close enough to
transfer energy each other, fluorescence light from one molecule can excite the other
molecule. It is called fluorescence resonance energy transfer (more commonly referred to
by the acronym FRET) (Herman, 1998; Jameson et al., 1989; Mason, 1993; Pavesi and
Fauchet, 2008; Ploem and Tanke, 1987; Rost, 1992).
For absorption of the excitation light to be efficient, the wavelength of light should
correspond to the energy difference between the ground state and an excited state. Figure
2-1 shows this process schematically. In Figure 2-1, heavy and thin horizontal lines
represent electronic and vibrational energy levels respectively, and upward and
downward arrows represent absorption and emission of a photon of light respectively.
Here, the vertical axis represents the energy of various levels. Also, in the figure, the
fluorescence excitation and emission spectra of the fluorophore are shown. The
magnitude of the absorbance and emission is determined by the probability of the
corresponding transition. The wavelength corresponds to the energy of the absorption and
emission transitions. The wavelength (λ) is related to energy (E) and frequency (ν) by the
equations:
where h is Plank’s constant and c is the speed of light (Guilbault, 1973; Morrison, 2008).
The absorption and emission spectra of fluorescence is determined by the energy distance
between the ground state and the electronically excited state. Here, when the emitted
photon has less energy than the absorbed photon, the energy difference between the
E = hv = hc /λ
12
longest wavelength absorbance maximum and the shortest wavelength emission
maximum is called the Stoke’s shift. If the emitted photon has more energy, the energy
difference is called an anti-Stokes shift (Lakowicz, 2006).
S0
S1
S2
S1
S0
Wavelength (nm)
Intensity
Energy
0
1 2 3
0
1 2 3
1 2 3
0
0 1 2 3
1 2 3
0
Excitation Emission
Figure 2-1. Jablonski diagram of excitation and emission, and the corresponding excitation and
emission spectra: It shows electronic and vibrational energy levels and absorbance and emission
transitions between energy levels for fluorophores.
In Table 2-1, the peak wavelengths of absorbance and emission of a variety of
fluorophores are shown.
Dye Absorbance
Wavelength(nm)
Emission
Wavelength(nm) Visible color
DAPI 345 455 Blue
Alexa fluor 488 494 517 green (light)
Fluorescein FITC 495 518 green (light)
Cy3 550 570 yellow
Rhodamine 123 560 580 yellow
Alexa fluor 568 578 603 red
Cy5 650 670 red
Alexa fluor 680 679 702 red
Cy7 743 770 red
Table 2-1. Peak excitation and emission wavelength of fluorophores.
13
Among the fluorophores, those with near infrared excitation and emission wavelength
within optical window (600nm-1100nm) are more useful for small animal imaging than
other fluorophores with blue or green excitation and emission wavelengths since the
infrared light can penetrate deeper through tissues than green or blue light (light with
longer wavelength produces less scattering in tissue and the absorption coefficients of
tissues within the wavelengths is lower).
2.1.2. Intrinsic parameters of fluorescence
Fluorescence of fluorophores has several intrinsic parameters: fluorescence intensity,
excitation spectra, emission spectra, and lifetime. The parameters are widely used in
optical imaging. The fluorescence intensity depends on quantum yield (Φ) and the
concentration of fluorophores. The quantum yield displays the efficiency of the
fluorescence process. The definition of it is the ratio of the number of photons emitted
(Npe) to the number of photons absorbed (Npa) in fluorophores.
Φ = Npe / Npa
Here, maximum quantum yield is 1.0. The quantum yield of 0.1 is still considered quiet
bright(Morrison, 2008).
Fluorescence lifetime is also a key parameter of fluorescence that is considerably
sensitive to their environment. The fluorescence lifetime can indicate the average time
that the molecule stays in its excited state before emission. An excited fluororphore will
return to the ground state with a certain probability based on the decay rates through a
14
number of radiative and/or nonradiative decay processes. It follows the first-order rate
equation:
* * * /
0 0
F = F e−kFt = F e−t τ F
Where *F is the number of electronically excited fluorophores at certain time t and *F0 is
initial fluorescence at t=0, kF is the rate constant by both radiative and nonradiative
processes and τF is the lifetime of the excited state, equal to 1/ kF, and therefore equal to
the fluorescence lifetime (Valeur, 2002).
The fluorescence lifetime and intensity can be affected by photon quenching. There are
collisional quenching and static quenching. Collisional quenching involves collisions
with other molecules that result in the loss of excitation energy as heat instead of as
emitted light. This process is always present to some extent in solution samples; species
that are particularly efficient in inducing the process are referred to as collisional
quenchers (e.g. iodide ions, molecular oxygen, nitroxide radical). Static quenching is
caused by interaction of the fluorophore with the quencher, which forms a stable non-fluorescent
complex. Since this complex typically has a different absorption spectrum
from the fluorophore, presence of an absorption change is diagnostic of this type of
quenching (by comparison, collisional quenching is a transient excited state interaction
and so does not affect the absorption spectrum). A special case of static quenching is self-quenching,
where fluorophore and quencher are the same species. Self-quenching is
particularly evident in concentrated solutions of tracer dyes (Komiyama and Miwa, 1980;
Lakowicz, 1999; Merkle et al., 1987; Prasad et al., 1983).
15
2.2. Optical imaging techniques for small animal imaging
2.2.1. Single-wavelength, single-photon excited fluorescence intensity imaging
Traditional fluorescence imaging has been a simple method to continuously track
movements and concentrations of labeled molecules in vivo, based on the spatial
distribution of intensity at a single emission wavelength. This technology allows
monitoring the progression of diseases, the effects of drug candidates on the target
pathology, the pharmacokinetic behavior of drug candidates, and the treatment outcomes
(Bouvet et al., 2002; Chen et al., 2004; Contag, 2007; Contag and Bachmann, 2002;
Farkas and Becker, 2001; Hoffman, 2004, 2005; Paulmurugan et al., 2004; Rice et al.,
2001; Vooijs et al., 2002; Yamauchi et al., 2005; Yang et al., 2000). It is based on the
linear dependence of fluorescence intensity on the accumulated concentration of labeled
molecules.
Figure 2-2. Fluorescence intensity image: (A) Fluorescence intensity image of a mouse with a
brain tumor. After the injection of tumor-targeted drug molecules stained with Alexafluor 680,
the image was obtained with 680nm excitation light and 710nm emission filter. (B) Absorption
and emission spectra of Alexafluor 680.
(A) (B)
Brain tumor
Excitation
Emission
16
Figure 2-2 shows the fluorescence intensity image of a nude mouse. In the figure, we can
see drug molecules stained with Alexafluor 680 are more accumulated in brain tumor
regions than other regions. However, it is very difficult to discriminate and quantify the
fluorescence signal of interest in the presence of confounding signals (noise) from entities
such as endogenous fluorophores. Thus, for more accurate quantitative imaging in the
presence of such “noise”, it is necessary to use more advanced imaging methods (Contag,
2007; Farkas and Becker, 2001).
2.2.2. Spectral detection
Spectral imaging provides a high-resolution spectral signature at every pixel of an image.
The spectral signatures are the specific combination of reflected or absorbed light at
varying wavelengths which can uniquely identify an object or a biological molecule
(Burton et al, 2009). This capability provides useful information compared to standard
fluorescence intensity imaging with a single excitation/emission channel. In spectral
imaging, a series of images of a sample through a liquid crystal tunable filter (LCTF) or
an acousto-optical tunable filter (AOTF), which only allow passage of light of a narrow,
chosen bandpass, are acquired in order to create a spectral data “cube” (x, y, and
wavelength) (Gebhart et al., 2007; George and Patonay, 1997; Gupta, 2005, 2009; Lee et
al., 2007; Zuzak et al., 2008; Zuzak et al., 2007). In this cube, spectral signatures can be
obtained at every pixel of an image. The spectral signatures are analyzed by various
spectral classification techniques such as Euclidean distance measure, correlation
measure, spectral angle measure, and spectral unmixing. Figure 2-3 shows the band
17
sequential image acquisition, spectral data cube, selected spectral signatures from the
cube, and the classified image for spectral imaging. In the classified image, red blood
cells (red pseudocolor) are quantitatively discriminated from other cellular matrix (green
color). Ultimately, the spectral “signatures” allow a very delicate, specific and
reproducible separation of signals of interest from noise, in these particular circumstances.
Also, additional information may be obtained by comparing spectra obtained at different
excitation wavelengths (Periasamy, 2001). The entire process can be completed in several
seconds (Burton et al., 2009; Carano et al., 2004; Chung et al., 2006; Farkas and Becker,
2001; Levenson et al., 2008; Levenson and Mansfield, 2006; MacKinnon et al., 2005;
Macville et al., 2001).
Figure 2-3. Spectral imaging: (A) A band sequential image acquisition using AOTF or LCTF. (B)
Spectral data cube. (C) Spectral signatures from selected pixels. (D) Spectral classification: a
spectral data cube is constructed through a (band-sequential) series of images acquisition along a
wavelength axis, with a narrow bandwidth. Each pixel contains its own spectral signature.
Spectral classification was performed using linear discriminant analysis methods.
(B) (C) (D)
(A)
Object
Collecting
Lens
AOTF,
LCTF
Detector
Wavelength
500 600 700 800
Intensity (A.U.)
0
1
18
2.2.3. Fluorescence lifetime detection
Fluorescence lifetime detection is a powerful tool to provide an image based on a
fluorophore’s fluorescence lifetime rather than intensity. The fluorescence lifetime is less
dependent on transient changes in the local concentration of the fluorophore but can be
highly sensitive to physiological environment such as intra/extracellular, pH distribution,
blood flow, tissue oxygen, and temperature (Bloch et al., 2005; Cubeddu et al., 1997;
Cubeddu et al., 1993; Esposito et al., 2006; Feng and Huang, 2007; Gratton et al., 2003;
Hanson et al., 2002; Lin et al., 2003; Lin et al., 1999; Murata et al., 2000; Niesner et al.,
2008; Spriet et al., 2008). Thus, unlike intensity or spectral measurements, fluorescence
lifetime imaging can show the energy transfer rate from the excited state of fluorophores
to its surrounding environment. Also, it can be ued for localizing diseased tissues such as
tumors (Bloch et al., 2005; Pelet et al., 2006).
Fluorescence lifetime detection can be performed either in the frequency-domain or time-domain.
Frequency-domain fluorescence lifetime detection is simpler and easier to
implement as well as have more light efficiency than time-domain fluorescence lifetime
detection. However, it requires more computational resources and has a limitation in
temporal dynamic range (Pelet et al., 2006). On the other hand, time-domain fluorescence
lifetime detection requires ultra-short pulsed laser but can provide better analysis of
complex exponential decays, particularly in the presence of several lifetime components
(Figure 2-4).
19
Figure 2-4. Fluorescence lifetime imaging: (A) Fluorescence intensity decay. After ultra-short
pulsed light excitation, fluorescence decays over time. (B) Fluorescence images over time. They
are obtained by ultra-fast time-gated camera. From these consecutive images, the fluorescence
intensity decay curve at every pixel can be generated. (C) Pseudo-color mapped image of
fluorescence lifetime. Each fluorophore has its inherent fluorescence lifetime.
In time-domain fluorescence lifetime detection, consecutive images are collected with
certain delay steps over time. Then, the decay function (intensity vs. delay time) is
measured at every pixel of the images and the measured decay curve is fitted to the first
order exponential decay curve in order to acquire time constants (lifetime) of the decay
curve. Finally, the lifetime obtained at every pixel is displayed with a color map
(Cubeddu et al., 1993). Figure 2-4 shows the fluorescence lifetime imaging of H&E
stained breast tissue. In this dissertation, time-domain fluorescence lifetime detection was
mainly used for the experiments.
2.2.4. Two-photon excitation
Two-photon fluorescence excitation is generated by simultaneous absorption of two
photons with half the energy (of an equivalent single photon) for excitation. This
0.3ns
0.5ns
3ns
2ns
1.5ns
1ns
0.3ns
0.5ns
3ns
2ns
1.5ns
1ns
(A)
(B)
(C)
Time
(C)
Lifetime (ns)
2.0
0
Time
Excitation
Pulsed laser
e−t/τ 2
e−t/τ1
20
phenomenon requires high fluxes of photons within an ultra-short pulse width. As a result,
a confocal-type effect (i.e. 3-D spatial selection/slicing) is generated because excitation
occurs only in the confined volume around the plane of focus when using a microscope
objective with a high numerical aperture. Figure 2-5 shows the typical Jablonski diagram
and geometry of two-photon excitation. Figure 2-5(A) shows the two-photon absorption
in a molecule. Figure 2-5(C) shows the 3-D point spread function of two-photon
excitation (Diaspro and Robello, 2000; Dong et al., 2003; Ibanez-Lopez et al., 2005). In
addition, since near infrared (IR) light is used for two-photon excitation, less photo-bleaching,
less photo-toxicity, and a deeper penetration can be obtained. Because of these
advantages, scanning two-photon fluorescence microscopy has been widely used as a
noninvasive method in tissue and animal imaging (Chirico et al., 2003; Denk et al., 1990;
Patterson and Piston, 2000).
To name only a few examples, it enables measurement of calcium dynamics deep in brain
slices of live animals (Mainen et al., 1999). In cancer research, this technique has been
used for in vivo studies of angiogenesis and metastasis (Kirkpatrick et al., 2007; Konig et
al., 1996; Wessels et al., 2007). Also, video-rate multiphoton microscopy has been
recently developed and used for studies of lymphocyte motility and antigen response in
intact lymph nodes (Miller et al., 2002). Recent advances in multiphoton in vivo
microscopy include the development of multiphoton endoscopes based on GRIN lenses
(Jung and Schnitzer, 2003).
21
Figure 2-5. Jablonski diagram and geometry of two-photon excitation: (A) Jablonski diagram of
two-photon absorption/florescence: two photons are simultaneously absorbed in fluorophores for
excitation. (B) Geometry of two-photon excitation. the two photon excitation is generated by
ultrashort pulsed light, and it has intrinsic sectioning ability and provides less photobleaching. (C)
Point spread function of two photon excitation.
As previously mentioned, this two-photon excited fluorescence imaging is very useful for
in vivo tissue and live animal imaging. Thus, it will allow immediate high quality
pathologic assessment of intact unstained tissues or regions of interest in a live animal
with high spatial resolution. In this dissertation, either the scanning two-photon excited or
wide-field two-photon excited fluorescence imaging were used for the acquisition of high
resolution information from intact unstained tissues and small animals.
2.2.5. Intra-vital confocal
Intra-vital confocal imaging facilitates high-resolution studies of cellular and molecular
events in vivo (Falati et al., 2002). In particular, intra-vital confocal imaging is based on
laser scanning confocal technology through an optical fiber bundle (diameter: ~1.5mm)
[Figure 2-6(A)]. The optical fiber bundle, which is a high resolution endoscope, can be
inserted in a small animal. Thus, it can enable us to observe physiological processes as
well as various tissues of interest inside a small animal at the cellular and sub-cellular
levels.
(A) (B) (C)
22
Figure 2-6. Scanning method through a fiber bundle and the applications of the intra-vital
confocal mode (from Mauna Kea Technologies, Paris, France): (A) scanning schematic. (B)
Blood vessels. (C) a mouse subcutaneous tumor.
Currently, this technology has been utilized for various applications in vivo such as the
observation of a subcutaneous tumor, peripheral nerve imaging, and angiogenesis.
However, it has some limitations; the specificity is relatively low due to a fixed filter, and
artifacts can be easily introduced to an image as the detector is not steady during data
acquisition (Laemmel et al., 2004; St Croix et al., 2006).
2.2.6. Bioluminescence
Bioluminescence imaging enables the noninvasive study of biological processes without
excitation light sources in small animals. For in vivo bioluminescence imaging, a reporter
gene, which encodes light-generating enzymes (luciferase), labels biological entities such
as bacteria, tumor cells, immune cells, or genes. Luciferase enables generation of visible
light by oxidation of an enzyme-specific substrate (luciferin) in the presence of oxygen
and ATP, without any excitation.
(A) (B) (C)
23
Luciferyl adenylate Oxyluciferin
+ O2 +
Luciferase Light,AMP
Luciferin + ATP Luciferyl adenylate + PPi
Figure 2-7. Chemical reaction process of luciferase enzymes and enzyme-specific substrates
(luciferin): the luciferin becomes luciferyl adenylate in the presence of ATP as an energy source,
and then, light is generated by that luciferase catalyzes the oxidation of the luciferyl adenylate.
Right side of figure shows a crystal structure of firefly luciferase complexed with oxyluciferin
and adenosine monosphate (AMP).
Figure 2-7 shows the chemical reaction process of luciferase enzymes and enzyme-specific
substrate. Bioluminescence imaging thus has several advantages compared to
fluorescence imaging: (i) It provides higher sensitivity; (ii) It has the capability to
observe the emitted visible light from an enzymatic reaction at few centimeters depth;
(iii) Imaging background from excitation or animal autofluorescence can be dramatically
reduced because no excitation light source is needed after the injection of a luciferin to
initiate the reaction. Thus, the signal to noise ratio is increased; (iv) It can offer similar
capabilities to SPECT and PET, besides optical imaging (Bhaumik et al., 2004; Contag,
2007; Contag and Bachmann, 2002; Graves et al., 2003; Patsenker et al., 2008;
Paulmurugan et al., 2004).
2.3. Effects of the tissue window on both excitation and emission photons
Figure 2-8 shows optical characteristics of tissues. In the figure, the light absorption of
the hemoglobin significantly decreased over 600nm as well as the light absorption of
water is reduced below 1100nm. Thus, the wavelength range, approximately 600nm-
1100nm, is called an “optical window” (Parrish, 1981).
24
Figure 2-8. Optical window for tissue imaging.
2.4. Chemotherapy molecules for cancer treatments
2.4.1. Nanoconstructs
In chemotherapy for tumor treatment, targeted delivery enhances efficiency and tumor
specificity, and thus (by allowing lower doses) reduces systemic toxicity. A
nanoconstruct (Polycefin) was developed as a novel platform for tumor treatment. The
nanoconstruct can carry more than one drug, allowing synergistic effects on tumor cells.
Figure 2-9 shows a schematic of the nanoconstruct.
Figure 2-9. Schematic of Nanoconstructs (Polycefin) (Ljubimova et al., 2008).
Drug
releasing
unit
Inhibitor 1 Inhibitor 2
Tumor cell
targeting
antibody to
transferrin
receptor
Protector
(PEG)
Fluorescent
dye for drug
tracking
Endosomal
disruption unit
for drug release
inside the cell
Optical window
Hemoglobin
Water
Wavelength (nm)
Absorption
0 600 800 1000 1200 1400
0.001
0.01
0.1
1
10
25
It consists of several modules chemically conjugated to poly β-L-malic acid (PMLA).
Here, for the targeted delivery of antisense oligonucleotides (AON, targeting the mRNAs
of laminin-8 α4 and β1 chains) into certain tumors, antibodies (mouse anti-human TfR
mAb) are conjugated to the PMLA. In addition, several functional units are chemically
conjugated to the carboxyl groups of the PMLA platform. Polyethylene glycol (PEG)
protects against enzymatic degradation. A dye (such as Alexafluor 680) is conjugated for
in vivo fluorescence imaging. L-Leucine ethylester moieties function in endosomal
escape. This nanoconstruct is biodegradable, non-immunogenic, and non-toxic (Lee et al.,
2006; Ljubimova et al., 2008).
2.4.2. Targeted sulfonated gallium corroles
Recent inventions of chemotherapy molecules for cancer prompt the development of new
multifunctional technologies aimed at improved detection, diagnosis, and/or therapeutic
intervention of cancer (Kawasaki and Player, 2005). The 2,17-bis-sulfonated corrole and
its metal complexes have similar chemical structure with porphyrins and associated
macrocycles that have been explored for cancer therapy, but may have distinct
advantages over the latter compounds for physiological application. Specifically,
sulfonated corroles are water soluble, thus it enables facile use in physiological fluids.
Also, they are unable to enter cells without a carrier molecule, thus it enhances the
specificity of delivery. Furthermore, any excitation is not required in order to elicit
cytotoxicity (thus expanding the potential tissue depth and distance at which corrole-mediated
therapy may be administered).
26
Figure 2-10. Chemical structure of sulfonated gallium corroles (from Harry B. Gray).
Here, for breast cancer treatment and detection, a breast cancer-targeted carrier protein
(HerPBK10) that promotes uptake by the cell and a sulfonated gallium-metallated corrole
(S2Ga) are synthesized. The S2Ga forms a tight assembly with the carrier protein that
resists high speed centrifugation and transfer to albumin, brightly fluorescence, and
induces toxicity to target cells after delivery and uptake by the carrier. Importantly,
corrole cytotoxicity is best supported by a membrane penetrating function, as non-penetrating
carriers such as albumin did not enable sufficient cytotoxicity. This suggests
that sulfonated corroles entering cells via receptor-mediated endocytosis must somehow
escape the endosomal vesicle to induce cytotoxicity. The protein used in these studies
provides both the targeting and penetration required for effective corrole delivery
(Agadjanian et al., 2009; Agadjanian et al., 2006).
27
CHAPTER 3
METHODS AND MATERIALS
3.1. Multimode optical imaging system
A new multimode optical imaging system is designed to be application-optimizable, with
higher sensitivity and specificity. This instrument combines various in vivo imaging
modes, including one or two photon-excited fluorescence, spectral, fluorescence lifetime,
intra-vital confocal, and bioluminescence imaging.
3.1.1. Design
A schematic diagram of the instrument we designed and built is shown in Figure 3-1(A).
It consists of a light-tight enclosure within which various optical imaging modes are
implemented as follows: fluorescence intensity, spectral, lifetime, intravital confocal,
scanning/wide-field two-photon excited fluorescence, which has been developed outside
but will be incorporated into the system, and bioluminescence imaging. Excitation of the
molecules of interest is made with laser light at the appropriate wavelengths. A wide
selection of light sources is available, including a tunable femtosecond pulsed laser
(MaiTai, SpectraPhysics), a picosecond pulsed laser (Tsunami, SpectraPhysics),
continuous wave (CW) lasers, and a broadband ps pulsed light source covering the whole
visible and near infrared range (450-1900 nm, Fianium). The light sources are delivered
28
via free space, or through a fiber-based laser delivery system. The fiber-based laser
delivery system allows flexible delivery of various lasers inside a light-tight enclosure.
Figure 3-1. Multimode optical imaging system: (A) System schematic. This system is capable of
several imaging modes, including fluorescence intensity, spectral, life-time, intra-vital confocal,
and bioluminescence imaging. Also, for 3D fluorescence imaging, optical components and a
control program installed. Furthermore, scanning/WTEF imaging mode will be incorporated with
the system for deep tissue imaging at high resolution. (B) Photographic image of the multimode
optical imaging system.
(B)
Intra-vital confocal imaging
unit
Anesthesia machine
AOTF
FLIM
Scanning/widefield two-photon excited
fluorescence imaging unit
Laser delivery
(A)
CCD
Moving
Stage
Time-gated Joystick Controller
anesthesia
machine
Motorized
filter wheel
3D Fluorescence
Imaging
Fiber-based delivery
system/small opening
AOTF
FLIM
Diffuser
Output Wavelength
- 405nm Laser : 405nm
- PS Laser : 700~1080nm
- HeCd Laser : 442nm
- FS Laser : 710~990nm
- Ar-Kr+ Laser : 488nm,
514nm, 647nm
Laser - HeNe Laser : 633nm
Telecentric
Heating pad
(X, Y)
(R)
Femtosecond pulsed
Argon-Krypton
HeNe
405nm
Picosecond pulsed
Mirror
HeCd
Optical probe
for intravital confocal
imaging
29
The light source, which is delivered into the light tight enclosure, is divided to two
directions by a 50:50 dichroic mirror. And then, they are delivered onto specimens
through diffusers and mirrors. Here, uniform excitation of the specimen is realized by
scattering the laser light onto the 90% transmission broadband diffusers (THORLABS, 1”
round 200 circle tophat diffuser). Light from the specimen is collected by a telecentric
lens (MELLES GRIOT, InvaritarTM 59LGL428 and 59LGG950, NA: 0.24). The collected
light passes through either (i) standard interference filters (Chroma) installed in a
motorized filter wheel, (ii) an AOTF or (iii) a FLIM module (LAVISION, PicoStar HR)
before arriving onto a cooled charge-coupled device (CCD) camera (Hamamatsu, ER-ORCA
or Princeton Instruments, PIXIS 400) located on top of the light-tight imaging
chamber. Every component is modular, and switching between modes is fast (~seconds,
if necessary). The AOTF and FLIM modules can be flexibly exchanged by using a
sliding rail. In addition, an optical probe (Manua Kea Technologies) has been
incorporated for intravital confocal imaging in this system. Also, for 3D fluorescence
imaging, the telecentric lens and CCD camera can be attachable onto the enclosure’s side
wall. The synchronization between a rotational stage and CCD camera for image
acquisition with different angle views is controlled by a program [Figure 3-4(E)] we
developed. Furthermore, a two-photon excitation imaging module is developed for high
resolution deep tissue imaging.
The animal (usually a mouse or rat) is placed inside the light-tight enclosure on a moving
stage that has a spatial resolution of 5μm in x and y-axis, 100μm in z-axis, and 2.16
arcsec in rotational axis. The sample stage can move along x-, y-, z-axis, and rotate
30
clockwise and counterclockwise to vary the field of view and a sample position. The
stages are actuated by a stepper and a servo motor using software we developed (using
CVI/National Instruments). In order to prevent animals from moving during image
acquisition, a gated anesthesia system we developed (capable of stopping breathing for
the duration of image acquisition alone) that uses a mixture of oxygen and isoflurane is
attached to the imaging box.
Figure 3-2. Laser delivery system: (A) Laser beam steering schematic. There are two light paths.
While the first light path is for pulsed laser light, the second light path is for continuous laser
lights. For the prevention of damage of pulsed laser from undesired back propagation, a Faraday
rotator and a retarder are installed. (B) Spectrum of the laser light in the system.
􀂄􈐴 405nm diode laser (405nm, 25mW)
􀂄􈑈 HeCd laser (442nm, 18mW)
􀂄􈑁 Argon-Krypton laser (488nm, 514nm,
and 647nm
􀂄􈑈 HeNe laser (632.8nm,25mW)
􀂄􈑐 Picosecond pulsed laser (700~1080nm,
1.6W)
􀂄􈑆 Femtosecond pulsed laser (710~990nm,
1.8W)
􀂄􈐶 670nm diode laser (670nm, 150mW)
(inside)
(A)
405
405nm diode laser
441.6
Helium Cadmium
457.9
476.5
488
501.7
514
Argon ion laser
530.9 568.2
Krypton ion laser 632.8
He-Ne laser
Krypton ion laser
670
647.1
670nm diode
laser
Retarder
a light-tight enclosure
Flip mirror
Fixed Mirror
10.5"
HeCd laser
Fs Pulsed Laser
Ps Pulsed Laser
BBO
405nm Laser
Faraday
rotator
HeNe Laser
Argon Krypton Laser
External trigger
Wheel filter
Lens
95:5 Mirror
Pico-, Femto-second
pulsed laser
(B)
31
3.1.2. Laser delivery
For excitation of molecules of interest, various laser lines (deep blue diode, 670nm diode,
picosecond pulsed, femtosecond pulsed, HeCd, Argon-Krypton, and HeNe laser) can be
used like as in Figure 3-2 (with their respective excitation wavelengths).
Each laser light source is selected by flipping mirrors and is delivered onto a specimen
through mirrors. In particular, a Faraday rotator are used as an optical isolator in order to
prevent undesired back propagation of light from disrupting or damaging a femtosecond
and a picosecond pulsed laser. Also, a Barium-borate crystal (BBO) is used for a second
harmonic generation of near-infrared lights in fluorescence lifetime imaging.
3.1.3. Fluorescence light detection
A telecentric lens (Navitar) is utilized in order to collect fluorescence emitted from a
specimen. The telecentric lens provides no magnification distortions due to an object's
distance or position in the field of view.
The collected light passes through a selected filter installed in a rotational filter wheel,
and then are recorded in a CCD camera (Pixis 400 or Hammastu ORCA-ER) after
traveling through a relay lens. The relay lens allows introduction of a filter between a
telecentric lens and a CCD camera (Figure 3-3).
The CCD camera can be replaced by a spectral or a fluorescence lifetime imaging module
depending on applications. The detection characteristics of combination of a Pixis 400
32
camera and a telecentric lens have been examined. The resolution of image is 0.222mm,
the field of view 111mm x 88.8mm, the working distance 180mm, and a depth of focus
40mm.
Figure 3-3. Relay lens system for filter installation: we installed a relay lens between them in
order to introduce a rotational filter wheel, which includes emission filters, between the lens and
CCD. The relay lens has 19.2mm image/objective focal distance.
3.1.5. Control software
The stage control board is under the control of a program that we developed using
CVI/National Instruments. The program panel is shown in Figure 3-4. It allows us to
control every stage in the multi-mode optical system as well as a motorized filter wheel.
Also, the program can be connected with client programs for control of the system from
other locations (Joystick/FLIM), 3D fluorescence/bioluminescence, and scanning/wide-field
two-photon excited fluorescence imaging through an internet connection. Figure 3-5
describes overall system control schemes. A system control program is connected with a
stage, a filter wheel control board, and a data acquisition board through universal serial
bus (USB) port in a computer as well as it is connected with Joystick/FLIM, 3D
fluorescence/bioluminescence imaging, and scanning/Wide-field two-photon excited
f1=17.526mm
Rotational filter wheel
6.35mm
Telecentric lens
f2=19.2mm
f3=19.2mm
Relay lens
(19.2:image/objective distance)
CCD
Rotational filter wheel
Relay lens
(Image/objective
distance: 19.2mm)
f3=19.2mm
f2=19.2mm
f1=17.526mm
Telecentric lens
33
fluorescence imaging program through internet. Also, the stage and filter wheel control
board is connected with a x, y, z-axis translation, rotation stage, and a motorized filter
wheel. A data acquisition board (National Instruments) receives data from a
scanning/wide-field 2-photon imaging module, and transfers control signals to light-emitting
diode (LED).
Figure 3-4. System control program: (A) Positioning buttons for controlling x-, y-, z-axis
transition, and rotation stage. (B) Emission filter selection. It allows us to exchange emission
filters installed in the motorized filter wheel. (C) Scanning FLIM control panel. In this panel,
scanning area and scanning step size can be selected. (D) Scanning/wide-field two-photon excited
fluorescence imaging panel. (E) Connection panel for joystick and scanning FLIM. (F)
Connection panel for 3D fluorescence imaging.
Stage and filter wheel
control board
System control
program
x, y, z, and
rotational stage
Motorized
filter wheel
Joystick/FLIM
3D fluorescence/
bioluminescence
USB Internet imaging
Scanning/wide-field
2-photon imaging
D-sub15
Data acquisition
board
Figure 3-5. Overall system control scheme.
(E)
(D)
(F)
(A)
(B)
(C)
34
3.2. Imaging modes in a multimode optical imaging system
3.2.1. Single-wavelength, single-photon excitation fluorescence intensity imaging
In this mode, a laser source can be selected by flipping mirrors for the excitation of a
variety of fluorophores. The selected light source is delivered onto a specimen through
mirrors, filters, and diffusers. Before a fluorescence image from the specimen is acquired,
a background image including a spatial profile of an excitation light source is recorded
for flat-field correction. Then, fluorescence collected by a telecentric lens is recorded in a
high sensitive cooled CCD camera (Pixis 400), after being selected by an emission filter.
Finally, the fluorescence image is corrected by the background image in order to reduce
the artifacts due to the profile of an excitation light source. Sometimes, a photographic-equivalent
image is taken with LED illumination for overlaying a fluorescence image.
Figure 3-6 shows the schematic of the fluorescence imaging mode.
Figure 3-6. Fluorescence imaging mode: as mentioned above, six CW and two pulsed laser lines
can be selected for excitation of fluorophores. The selected light sources divided by a 50:50
dichroic mirror in a light-tight enclosure are delivered onto a specimen through two diffusers.
Then, fluorescence collected by a telecentric lens from a specimen, passed through selected
interference filters, is recorded on a CCD.
Lasers
Filter (bandpass
filter)
Mirror
Diffuser
Scanning
CCD
Computer
Filter
LED
Stage
controller
35
3.2.2. Spectral imaging mode
In this spectral imaging mode, a laser delivery for excitation of specimen is same with
that in fluorescence intensity imaging. In detection of fluorescence, emitted fluorescence
from a specimen is collected by a telecentric lens and passes through a long-pass filter
which rejects the excitation light source (Figure 3-7).
Figure 3-7. Spectral imaging mode: the excitation procedure is the same as the procedure of
single-wavelength, single-photon fluorescence intensity imaging. An acousto-optical tunable
filter is employed between a filter and CCD for band-sequential spectral selection. In the AOTF,
wavelength selection is realized by a high-frequency acoustical compression wave applied to a
crystal of tellurium dioxide or quartz in order to locally alter the refractive index of the crystal in
a periodic pattern. This leads, in effect, to the generation of a (virtual) diffraction grating so that
an orthogonal beam of polarized and collimated light incident at the Bragg angle is diffracted into
the first order beam, with selection of a particular wavelength, under computer control, by the RF
frequency applied.
Then, the light is delivered onto a high sensitivity, low noise CCD camera cooled to -
700C (Princeton Instruments, PIXIS 400) for imaging through an imaging AOTF system
we developed (ChromoDynamics, Inc.) which can rapidly and sequentially select a
narrow bandwidth (1.5-4.0nm). After the sequential images within the certain range of
wavelength are acquired, the spectral signatures on each pixel are generated by our
developed program (Chung et al., 2006). Then, the image is classified based on the
Lasers
Filter (bandpass
filter)
Mirror
Diffuser
Scanning
CCD
Computer
AOTF
Filter
Stage
controller
36
selected spectral signatures. In addition, spectral unmixing is performed in order to reject
autofluorescence from fluorescence signals of interest using Image J (Gammon et al.,
2006).
3.2.2.1. Spectral computation methods
In the spectral classification techniques, the Euclidean distance measure is the most
simple spectral similarity measure method. In this dissertation, a root sum of square error
method (RSSE) in the Euclidean distance measuring methods is usually used for the
spectral classification:
2
1
( ( ) )/( )
N
i i
i
RSSE p r m M m
=
= Σ − − −
Where N represents number of spectra, pi and ri refer to the intensity of ith spectrum of
the sample pixel signature and reference signature respectively, and m and M are the
minimum and maximum of RSSE. Also, in the Euclidean distance measuring method,
there is the sum of area difference method. The sum of area difference method considers
the possibilities of the different steps in the spectrum axis (Grahn and Geladi, 2007).
On the other hand, the correlation measure indicates the strength and direction of a linear
relationship between two signatures. For spectral classification, there are the spectral
correlation similarity and spectral similarity value method based on the correlation
measure. While the spectral correlation similarity method based on the correlation
measure uses the Pearson correlation coefficient as a similarity measure, the spectral
37
similarity value method is a combined measure of correlation similarity and the Euclidian
distance.
The spectral angle measure method, which calculates the angle between two spectra in
order to measure spectral similarity between substance signatures for material
identification, has been widely used in spectral image analysis.
Also, the linear spectral unmixing method is very useful to analyze mixed signals to a
pixel of images obtained in spectral imaging. The linear spectral unmixing method is
based on the assumption that the recorded signal T at every channel λ can be expressed as
a linear combination of the contributing fluorophores. In the consecutive channels
obtained through spectral imaging, the distribution of emission signal represents directly
the fluorophore emission spectrum and creates a spectral signature. With linear unmixing
using these spectral signatures as reference, even combined and mixed signals can be
clearly separated into the fluorophores that contribute to the total signal (Farkas et al.,
1998; Zimmermann, 2005).
Spectral imaging with ratiometric analysis methods have been reported for dynamic
monitoring of other cellular processes and kinetics such as pH oscillations,
photobleaching, and quenching kinetics. The method was based on a variant of the
spectral waveform cross-correlation analysis. In the method, the master reference
spectral library is constructed by composite spectra with varying ratios of component
spectra (Ramanujan et al., 2006).
38
3.2.3. Fluorescence lifetime imaging mode
For fluorescence lifetime imaging in the multi-mode optical imaging system, we
developed a mosaic fluorescence lifetime imaging system with a femtosecond (fs) pulsed
laser, for a large field of view. In this method, fs pulsed laser light is tuned to 400~480nm,
a repetition rate of 80MHz, generated by the second harmonic of fs tunable pulsed laser
with 800~960nm (Mai-Tai Ti-Sapphire laser, Spectra-Physics) in a BBO crystal. It was
used for the excitation of molecules such as corroles or fluorescein inside a specimen. An
ultra-fast time-gated camera (LaVision, PicoStar HR) was utilized for fluorescence
lifetime imaging. Figure 3-8 shows the schematic of the mosaic fluorescence lifetime
imaging set-up. The fs pulsed laser can be delivered through a small opening in a light-tight
enclosure.
Figure 3-8. Mosaic fluorescence lifetime imaging using fs pulsed laser: fs pulsed light was used
for the excitation of the molecules of interest. The light was generated by the second harmonic of
infrared fs pulsed light in BBO crystal following a Faraday rotator. It is delivered onto a
specimen through the small opening on the light-tight enclosure via a built-in diffuser. A
specimen was scanned by excitation light while an x-y translation motorized stage, which is
controlled by the custom software, is moving. Fluorescence life time imaging was done by an
ultra-fast time-gated camera. The camera was synchronized with the laser light by an external
trigger and a delay unit. The image acquisition was performed by using a FLIM control program
which is connected with the actuation motor control program through a TCP/IP internet
connection.
Motor control program (Server)
Femto-second laser
(100fs,80Mhz)
BBO
Filter
(<500nm)
Mirror
Diffuser
Scanning
Trigger
Delay unit
Computer
External
Time Gated Intensifier
(LaVison Picostar HR)
CCD
Computer
Internet Connection
Motor control
program (Client)
39
And then, the light passes through two mirrors and a diffuser in order to excite a
specimen. The beam size of the light is controlled by distance between a specimen and
the diffuser which makes a 20 degree divergence of light. The fluorescence light from a
specimen is collected by the telecentric lens. The fluorescence light passes through an
emission filter, and then is delivered onto the CCD connected with time gated intensifier.
The CCD and time-gated intensifier are synchronized with the pulsed light via a delay
unit which connects with the external trigger. Scanning for a large field of view is
performed sequentially after each image acquisition by x-y translation motorized moving
stage under the control of the program which has been developed using National
Instruments/CVI. It can control scanning width, height, and step size. Also, the program
can be connected with a Joystick program which allows us to control the motorized
stages and scanning at different locations through TCP/IP connection. Figure 3-8 shows
the mosaic fluorescence lifetime imaging system.
3.2.4. Intra-vital confocal imaging
Intra-vital confocal imaging enables in vivo endoscopic imaging with high resolution. In
this mode, a holder for positioning a fiber bundle probe has been constructed in order to
reduce the artifacts due to operators’ movements. The fiber bundle probe is directly
connected with a confocal scanner including a scanning unit, an excitation laser source,
an emission filter, and a detection module (an avalanche photodiode). Figure 3-9 shows
the schematic of intra-vital confocal imaging mode.
40
Figure 3-9. Intra-vital confocal imaging mode: S-probe with a diameter, 1.5mm is utilized for
intra-vital confocal imaging. The probe has 0μm working distance, 5μm lateral resolution, and
15μm axial resolution. The excitation and emission lights are delivered and detected through the
distal end of the probe. Then the confocal scanner generates a confocal image. The image
obtained using the probe has a 599x500μm field of view and 450x384pixels.
3.2.5. Bioluminescence imaging mode
A bioluminescence imaging mode was developed in order to detect ATP and enzymatic
activity of engineered nude mice. The experimental setup is shown in the Figure 3-10. A
photographic image of the mouse was recorded, using two diodes. Bioluminescence was
then collected by a telecentric lens (Melles-Griot, InvaritarTM) and imaged onto a cooled
CCD (Princeton Instruments, PIXIS 400). The bioluminescence image was threshold and
overlapped with the photographic image.
Figure 3-10. Bioluminescence imaging mode: After injecting enzyme-specific substrates
(luciferin) into an animal, the bioluminescence signal is collected by a telecentric lens, and then
recorded onto a cooled CCD. After bioluminescence imaging is completed, the photographic
image is acquired with LED illumination for overlaying the bioluminescence image. LED light
intensity is controlled by a program we developed.
Scannin
Fiber bundle
Confocal scanner
Probe positioning holder
Stage
controller
Computer
Telecentric lens
Scanning
CCD
Computer
Filter
LED
Bioluminescence
Stage
controller
41
3.2.6. Scanning/wide-field two-photon excited fluorescence imaging mode
Scanning/wide-field two-photon excited fluorescence imaging mode was developed in
order to examine the details of molecules or tissues at deeper locations in multimode
optical imaging in vivo, outside the system. For wide-field non-linear excitation, a
polarized fs pulsed laser (MaiTai – Spectra Physics) with 780~990nm, 50-300mW
average power, and 80MHz repetition rate has been utilized. The fs pulsed laser beam
passes through a Faraday rotator, several mirrors, and Galvo mirrors. Then, it is finally
focused on the back focal plane of an objective (Nikon 40x, 1.3 NA, oil, Nikon 40x, 0.75
NA, air, or Nikon 60x, 0.75 NA) through a doublet lens (L1) (Melles Griot, FL200) in
order to obtain a more-or-less parallel beam which makes a larger spot size at the sample.
The fluorescence from a specimen is collected by the same objective, and then passes
through short/band-pass filters and a tube lens (L2) before being recorded on a cooled
CCD. In addition, in order to compensate a non-uniform excitation generated by the
Gaussian profile of the beam, flat-field correction of a wide-field two-photon excited
fluorescence image was performed by a normalized and inversed Gaussian mask. The
Gaussian mask has been constructed through the convolution of the original image with a
Gaussian function (radius: over 40) at a focal plane using ImageJ. The corrected image, R,
is obtained by the entry-by-entry product of the original image O and the Gaussian mask
M (R=O.× M).
Moreover, for two-photon excited fluorescence imaging of in-vitro specimens, a scanning
unit is incorporated into the system. In the scanning two-photon imaging mode, the CCD
is replaced with a photomultiplier tube (PMT) (Hamamatsu, H6780-20), and the doublet
42
lens (L1) is removed from an optical beam path. Here, Galvo mirror and PMT output
ignals are synchronized by the program we developed. Figure 3-11 shows the
experimental setup.
L1
Objective
Specimen
Dichroic mirror
CCD L2
PMT
Galvomirror
fs laser
F2 F1
FR
Figure 3-11. Scanning/wide-field two-photon imaging modes: L1 is a doublet lens which enables
wider excitation and is removed in scanning two-photon imaging mode. F1 and F2 is a shortpass
filter (<700nm) and a bandpass interference filter, respectively. L2 is a tube lens.
3.3. Wide-field two-photon excitation for multimode optical imaging in vivo
The two-photon excitation can be combined with fluorescence spectral or lifetime
imaging for various applications. Currently, spectral or lifetime imaging has been utilized
in order to distinguish and characterize fluorophores targeted to different molecules in
intact cells and tissues as an analytical tool with the power of object visualization in
biomedical imaging and research (Deniset-Besseau et al., 2007; Macville et al., 2001). In
particular, the spectral imaging has been used for multicolor spectral karyotyping of
human chromosomes, spectral pathology and diagnosis (Schrock et al., 1996), functional
imaging of brain oxygenation (Shonat et al., 1998), and evaluation of melanoma in vivo
(Tomatis et al., 2005). On the other hand, fluorescence lifetime imaging has also been
43
used to discriminate fluorophores as well as identify the functional status around the
fluorophores since fluorescence lifetime is sensitive to the surroundings around
fluorophores such as pH distribution, blood flow, tissue oxygen, and temperature
(Hanson et al., 2002; Ramanujan et al., 2005). Moreover, in in vivo imaging, the
combination of two-photon excitation with the spectral or fluorescence lifetime imaging
can enhance capabilities to discriminate between multiple targeted molecules at the
deeper location compared to one-photon spectral and fluorescence lifetime imaging
(Deniset-Besseau et al., 2007; Lakowicz, 1996).
However, when a scanning two-photon excitation method is utilized with non-scanning
imaging device, such as acousto-optical tunable filter and ultrafast time-gated CCD
camera for spectral/fluorescence lifetime detection, it is difficult to use scanning and non-scanning
element together in good fashion. It may require complex and sophisticated
synchronization devices in combining them. Thus, this approach is too burdensome. In
addition, scanning two-photon excited fluorescence imaging is not fit for monitoring fast
processes with high magnification in vivo, since it is a relatively slow method despite the
numerous theoretical advantages in small animal imaging. Moreover, a motion artifact
can be easily introduced during scanning two-photon excited fluorescence imaging. Thus,
these limitations of scanning two-photon excited fluorescence imaging prompted us to
develop another approach, which allows the combination of two-photon excitation with
other non-scanning imaging methods in multi-mode optical imaging for more versatile
and fast in vivo imaging.
44
Recently, two-photon excitation as a video rate imaging method has been reported. This
report shows that the Gaussian excitation profile in the lateral plane and the lack of
inherent sectioning ability of a scanning multiphoton microscope does not allow three-dimensional
rendering of sectioned data (Fittinghoff et al., 2000). However, despite the
disadvantages of wide-field two-photon excited fluorescence imaging compared to
scanning two-photon excited fluorescence imaging, it might nonetheless be very useful in
small animal imaging in vivo and, furthermore, in multi-mode optical imaging using non-scanning
devices, which has the capability to provide quantitative and functional
information simultaneously.
3.3.1. Sample preparation
In order to examine inherent optical characteristics of wide-field two-photon excited
fluorescence imaging, a 16μm cryostat section of mouse intestine (specifically, the
filamentous actin prevalent in the brush border) stained with Alexa Fluor 568 phalloidin
(FluoCells prepared slide #4 (F-24631), Invitrogen) and 50μM gallium corrole solution
(Agadjanian et al., 2006) were prepared. Also, an eye specimen of an alzheimer’s mouse
was prepared for in vitro test of wide-field two-photon excited fluorescence imaging. In
addition, a nude mouse with implanted breast tumors was used for wide-field two-photon
excited fluorescence imaging in vivo as well as for spectral and fluorescence life-time
imaging with wide-field two-photon excitation. Furthermore, 100μM gallium corroles
and microspheres (Constellation fluorescent microsphere, Invitreogen) were diluted with
phosphate buffer saline solution (PH 7.4) as fluorophores for the in vivo experiments.
45
3.3.2. Experimental Setup for wide-field two-photon excitation
The experiment setup is almost similar to the setup we described in Figure 3-11. Here,
the fluorescence from a specimen, collected by the same objective, passes through a
dichroic mirror and either band-pass filters (Chroma Technology, 560nm ±8nm or
630nm±60nm), a band-sequential spectral selection device (acousto-optical tunable filter,
ChromoDynamics Inc.), or a time-gated intensifier (LaVision Picostar) before being
recorded on a cooled CCD for the combination of wide-field two-photon excitation and
multimode instrument.
Figure 3-12. Schematics of the experimental set-up for wide-field two-photon excitation: (A)
Multimode optical imaging with wide-field two-photon excitation: For spectral/lifetime imaging,
we attached an imaging AOTF and LAVISON fluorescence lifetime module onto a NIKON
microscope with multiple ports. Also, fs pulsed light was delivered onto a specimen through
several mirrors, a lens (L1), and an objective. (B) wide-field two-photon excitation set-up in a
scanning confocal imaging microscope: For comparisons among optical characteristics of one-photon,
scanning two-photon, and wide-field two-photon excited fluorescence imaging, this
experimental set-up was constructed in a scanning confocal microscope. For scanning two-photon
excited fluorescence imaging, fs pulsed light source was delivered by Galvo mirrors onto
the back focal plane of the objective (PlanApo, 40x/0.75) through a scanning confocal imaging
unit.
For one-photon excitation, Hg lamps incorporated with a microscopy (Eclipse TE2000,
NIKON) was utilized through excitation filters. Figure 3-12(A) shows the experimental
(A) (B)
fs laser FR
L1
Objective
Specimen
Dichroic mirror
CCD
Confocal scanning
Filter
fs laser FR
L
Objective
Specimen
Dichroic mirror
Mirror
AOTF
CCD
CCD TGI
46
set-up for wide-field two-photon excitation. Also, the set-up of wide-field two-photon
excited fluorescence imaging previously mentioned was combined with a scanning
confocal imaging system [Figure 3-12(B)] (Olympus IX70, Fluoview) in order to
compare the optical characteristics of one-photon, scanning two-photon, and wide-field
two-photon excited fluorescence imaging.
3.3.3. Flat-field correction and spectral/fluorescence lifetime Analysis
Flat-field correction of wide-field two-photon excited fluorescence images was described
previously. All flat-field correction images have been obtained through using the method.
In spectral imaging, spectral classification algorithms based on Euclidean distance
measure (RSSE) was utilized for spectral analysis which is performed by our own custom
software (Chung et al., 2006). Also, for fluorescence lifetime analysis, a single
exponential decay fitting method was used to fit the data using our developed analysis
program, which provides the time-constant of a single exponential curve fitted to data on
same pixel in a series of images and generates a fluorescence lifetime image mapped by
the time-constants on each pixel (Foss, 1970).
3.4. Experimental setup and materials for investigation of optical characteristics of S2Ga
3.4.1. Sample preparation
50μM 125μM, 250μM, and 500μM S2Ga solutions diluted with phosphate buffered
saline (PBS) (pH 7.5) were prepared in order to investigate concentration (in)dependence
47
of fluorescence lifetime of the corroles. In addition, 50μM S2Ga with different pH values
(5.0, 5.5, 6.0, 6.5, 7.0, and 7.5) were prepared for examination of pH dependence of
corrole fluorescence lifetime.
3.4.2. Experimental setup
For fluorescence lifetime imaging, fs pulsed laser light tuned to 424nm, a repetition rate
of 80MHz, generated by the second harmonic of fs pulsed laser at 848nm (Mai-Tai Ti-
Sapphire laser, Spectra-Physics) in a BBO crystal, was used for the excitation of the
sulfonated gallium corroles. An ultra-fast time-gated camera (LaVision, PicoStar HR)
was utilized for fluorescence lifetime imaging. Figure 3-13 shows the schematic of the
experiment set-up.
Figure 3-13. Experimental setup for corrole fluorescence lifetime imaging.
The 424nm fs pulsed laser is delivered to Nikon Microscope through a Faraday rotator,
macro lenses, the Barium-Borate crystal, a band-pass filter (425±5nm), and several
mirrors. The light then passes through a diffuser and is reflected to the back focal plane
fs pulsed laser
Mirror
Faraday rotator
Lens
BBO
Filter
Diffuser
Dichroic
mirror
Objective
ΔT chamber
Filters
TGI
CCD
48
of 40x objective (Nikon 40x plan fluor, NA: 0.75) by a dichroic mirror inside a filter cube
for the excitation of the corroles inside a ΔT microscopic imaging chamber. The
fluorescence from the corroles is collected by the objective and delivered onto the CCD
connected with time gated intensifier (TGI) through emission filters (625±250nm and
620±60nm). The CCD and time-gated intensifier are synchronized with the pulsed light
via a delay unit that connects with the external trigger. In related experiments, we used
the ΔT culture dish system (Bioptechs) in order to control the temperature of solution
inside the chamber to within 0.1 degrees, for corrole and live-cell imaging.
3.4.3. Fluorescence lifetime of S2Ga
For examination of the effect of concentration on fluorescence lifetime of the sulfonated
gallium corroles, 100μl corrole solutions with different concentrations were added into
the ΔT chamber, respectively. Then, fluorescence lifetime of these corroles at the
respective concentrations (50μM 125μM, 250μM, and 500μM) was measured at room
temperature (24.5ºC).
In addition, fluorescence lifetime of the 50μM corrole solutions with different pH (5.0,
5.5, 6.0, 6.5, 7.0, and 7.5) was measured at room temperature and 36.5ºC respectively.
While the measurements, the temperature was strictly (~0.1ºC) controlled by a ΔT culture
dish system. Also, before adding the corroles into the ΔT chamber, the pH of the corroles
was confirmed with a pH meter.
49
Finally, the effect of temperature on fluorescence lifetimes of corroles has been
investigated. Each corrole solution with different pH was added into the ΔT chamber, and
then the fluorescence lifetime of the corroles was measured at a room temperature
(24.5ºC), 36.5ºC, and again room temperature (24.5ºC) sequentially. Before measurement
of fluorescence lifetimes at the room temperature following the fluorescence lifetime
imaging at 36.5ºC, we waited for 20 minutes in order to cool down the temperature of
corrole.
3.5. Fluorescence lifetime imaging of HerGa and S2Ga on human cancer cells
Human breast cancer cells (MDA-MB-435) and U251 glioma cells were cultured in a
Bioptechs ΔT chamber for 24 hours in order to examine the fluorescence lifetime changes
of HerGa and S2Ga due to the cancer cell endocytosis. Also, 150 μM HerGa solutions
were prepared (Agadjanian et al., 2009).
After culturing the cancer cells in a ΔT chamber for 24 hours, fluorescence lifetime of the
corroles bound to the cells was measured at different time points over 60 minutes. Before
adding 200μl of 50μM corroles with pH 7.5 into the ΔT chamber, the medium inside the
chamber was removed, and then the chamber was washed with PBS (pH 7.4). During 60
minutes, temperature of corroles was maintained around 36.5ºC.
50
3.5.1. Data acquisition and Analysis
For the fluorescence lifetime analysis, a total of 25 images were acquired within 4800ps
with a 200ps time step, and a gate width of 600ps after excitation of corroles. The images
were analyzed using the first order exponential decay fitting method incorporated into the
DaVis software (LaVision) and our program. (We have more advanced methods available,
but these were not needed).
3.6. Measurement of mitochondria membrane potential variations of the breast cancer
cells induced by HerGa
Mitochondria in cell physiology and pathology plays a variety of roles, including
production of ATP, regulation of the cell redox state, participation in ion homeostasis,
transport of metabolites, lipid and amino acid metabolism, and cell death. These
important functions highly depend on mitochondrial membrane potential. Thus, the
simultaneous measurement of the mitochondrial membrane potential may offer insight
into both the basic energy metabolism and its dysfunction in pathologic cells (Solaini et
al., 2007).
Measurement of mitochondrial membrane potential utilizes fluorescent potentiometric
probes. The mitochondrial membrane potential in situ is mainly monitored by measuring
the accumulation of cationic fluorescent probes in response to the mitochondrial
membrane potential in living cells. Cationic fluorescent probes distribution across the
membrane is governed primarily by the Nernst equation (Farkas et al., 1989). Currently,
51
Rhodamine-123, Tetramethyl rhodamine methyl ester (TMRM), and tetramethyl
rhodamine ethyl ester (TMRE) have been used as cationic fluorescent probes for the
measurement of mitochondrial membrane potential. Among them, TMRM is best since
TMRM reduces mitochondrial respiration to a much lesser extent than rhodamine 123 or
TMRE (Gottlieb and Granville, 2002).
In this thesis, mitochondrial membrane potential variations of breast cancer cells are
monitored in real time using TMRM in order to investigate cell mechanism of HerGa.
For the experiment, 20nM TMRM solution was diluted with PBS 7.4. MDA-MB-435
breast cancer cells were cultured onto a 20mm cover slip for 24 hours. A NIKON
TE2000E microscope was to image cells. TMRM and HerGa fluorescence emission was
selected with a Cy3 filter (exitation: 545±30nm and emission: 610±75nm) and HerGa
filter set (excitation: 420nm±40nm and emission: 620nm±60nm). Also, a highly sensitive
CCD camera (Photometrics, Cool snap HQ2) was utilized to record an image.
3.7. Spectral imaging with a ratiometric analysis method for dynamic monitoring of
HerGa accumulation and clearance
3.7.1. Experimental setup
For the examination of HerGa accumulation and clearance, we utilized the spectral
imaging mode in our multimode optical imaging system shown in Figure 3-7. For the
excitation of HerGa injected into the nude mouse, fs pulsed laser light tuned to 424nm at
a repetition rate of 80MHz, was generated by the second harmonic of a fs pulsed laser at
52
848nm (Mai-Tai Ti-Sapphire laser, Spectra-Physics) and a 500nm longpass filter was
used to reject the excitation light.
3.7.2. Construction of spectral signatures
For the spectral classification, a total of 10 sequential images were recorded within the
spectral range of 500nm to 680nm with a step size of 20nm and a bandwidth of 10nm.
Spectral signatures on each pixel were generated by our custom software. Four different
reference spectral signatures with varying ratios of autofluorescence and HerGa
fluorescence spectra were constructed. First of all, we acquired the spectral signatures of
autofluorescence and HerGa fluorescence from sequential images of a mouse before and
after HerGa intra-tumor injection. While the spectral signature of autofluorescence (A)
has a peak at 500nm and decreases as wavelength increases, the spectral signature of
HerGa (H) has a peak at 620nm.
Figure 3-14. Composite spectra with varying ratios of spectra of autofluorescence and HerGa
fluorescence.
500 550 600 650 700
-100
0
100
200
300
400
500
600
700
Wavelength(nm)
Fl. Intensity (A.U)
500 550 600 650 700
-100
0
100
200
300
400
500
600
700
Wavelength(nm)
Fl. Intensity (A.U)
500 550 600 650 700
-100
0
100
200
300
400
500
600
700
Wavelength(nm)
Fl. Intensity (A.U)
500 550 600 650 700
-100
0
100
200
300
400
500
600
700
Wavelength(nm)
Fl. Intensity (A.U)
Autofluorescence HerGa
(1 : 0) (1 : 1) (1 : 2) (0 : 1)
53
The fluorescence from a nude mouse after IV injection of HerGa can be represented by
the linear superposition of the two spectral signatures (A and H) as:
S=a1A+ a2H
Where S is the composite spectral signature and an is the relative concentration of
molecules. We constructed four spectral signatures with 1:0, 1:1, 1:2, and 0:1 ratios of A
and H for spectral classification. The ratio indicates relative HerGa contribution to
autofluorescence. While the 1:0 ratio of intensities represents pure autofluorescence, the
0:1 ratio represents high relative contribution of HerGa to autofluorescence. And, the 1:1
and 1:2 ratios represent the transition between the pure autofluorescence and the large
amount of HerGa accumulation, respectively (Figure 3-14). Thus, this approach can give
an accurate quantitative estimate of the relative concentration of two set of fluorophores
and their variation in space and time in vivo (farkas et al., 1998; Burton et al., 2009).
54
CHAPTER 4
RESULTS
As mentioned above, our multimode optical imaging system can be operated in at least
six different image acquisition modes, all shown to acquire complementary information.
While excessive acquisition times prevent the use of all imaging modes concurrently, it is
envisioned that a specific number of modes and image collection sequences can be
developed that maximizes the useful information extraction for all application. In this
chapter, the evaluation of each imaging mode in our system is performed with in vivo and
in vitro specimens. The feasibility of multimode optical imaging in vivo from our system
is evaluated. Finally, Multimode optical imaging in chemotherapy assessment and cancer
detection are performed.
4.1. Single modality experiments
4.1.1. Single-wavelength, single-photon excitation fluorescence intensity imaging for
discrimination between fluorophores conjugated with a drug molecule
Single-wavelength, single-photon fluorescence intensity imaging (it will called
fluorescence intensity imaging later) has been performed using two laser light-sources
and two emission filters in order to discriminate rhodamine B labeled liposomes
containing Alexafluor 680 labeled proteins. 514nm and 647nm light sources from
55
Argon/Krypton laser have been utilized for the excitation of rhodamine B and Alexafluor
680, respectively. Also, 590nm±60nm and 700±15nm emission filters have been utilized
for the detection of fluorescence from the molecules. After the molecule solutions with
different concentrations (20μM and 100μM) were injected into two implanted tumor
regions respectively, fluorescence from rhodamine B and Alexafluor 680 was detected,
sequentially. Figure 4-1 shows the fluorescence intensity images. In the figure, although
514nm light source for the excitation of rhodamin B generates more autofluorescence
than 614nm light excitation for Alexafluor 680, the fluorescence of rhodamine B can be
easily distinguished from autofluorescence from the mouse. The average light intensity
around the regions also has been measured in order to examine a concentration
dependency. The average intensity over tumor regions at the right side is five times
higher than at the left side. This ratio is almost identical to the concentration ratio of
injection molecules.
Figure 4-1. Fluorescence intensity images of a mouse injected with drug molecules tagged by
rhodamine B and Alexafluor 680: (A) Fluorescence image by rhodamine B (B) Fluorescence
image of Alexa680 (C) The overlaid image of a rhodamine B, an Alexafluor 680 and a
photographic mouse image.
(A) (B) (C)
56
In order to quantitatively evaluate the quality of the fluorescence images recorded, we
compared the contrast of tumor areas in the images of a mouse obtained with a
commercial system and our system. Here, we injected tumor targeting drug molecules
into intra-vein of the mouse, and then the images were acquired after 24 hours. Figure 4-2
shows the comparison of the contrasts as calculated with average intensities within the
two circles shown. The contrast values calculated based on Michelson contrast
(Michelson, 1927) for the commercial system and our system are 0.12 and 0.26
respectively, with our system showing a roughly two-fold improvement over the best
commercially available preclinical imaging system (Figure 4-2). This is by no means the
highest ratio obtained, only a representative one.
Figure 4-2. Contrast comparison between images recorded with a commercial system and our
multimode optical system: (A) Intensity image obtained by the commercial system. (B) Intensity
image obtained by our multimode optical system. The contrast values were calculated based on
[It(Tumor)-Ib(background)]/(Ib +Ib).
4.1.2. Spectral imaging with spectral unmixing for quantification of fluorescein
We have extensive experience in preclinical and clinical spectral imaging (Farkas et al.,
1998; Burton et al., 2009; Frykman et al., 2008). The feasibility of spectral imaging in
our new multimode system has been investigated after 10μM fluorescein, diluted with
(A) (B)
57
PBS (pH 7.4), was injected into breast tumors implanted in the back of a nude mouse.
488nm laser light was utilized for the excitation of fluorescein inside the mouse, and then
a total of 12 images were recorded within the spectral range of 500nm to 720nm with a
step size of 20nm. The images were analyzed using software we developed. Spectral
classification based on a least mean square error method and spectral unmixing was
performed for separating fluorescein fluorescence from autofluorescence respectively.
Figure 4-3 shows the classified images obtained by spectral classification and spectral
unmixing. In Figure 4-3(A), although fluorescence of fluorescein around tumor regions is
shown, the right side of brightest regions (selected region by a red mark) displays similar
intensity with autofluorescence around the selected region (by a green mark). However,
the fluorescence of fluorescein is clearly discriminated from autofluorescence in the
spectrally classified and the spectral unmixed image shown in Figures 4-3(C) and (D).
Moreover, Figure 4-3(B) displays the difference of the spectral signatures of fluorescence
and autofluorescence.
Figure 4-3. Spectral imaging of a nude mouse with injection of fluorescein into implanted tumors:
(A) Fluorescence intensity image at 540nm. (B) Spectral signatures within the selected regions.
While the spectral signature of fluorescence has a peak at 540nm, the spectral signature of
autofluorescence has multiple peaks and shows relatively broader spread of wavelength. (C)
Spectral classification image by the selected signatures. While red color shows fluorescence of
fluorescein, green color represents autofluorescence. (D) Spectrally unmixed image. After
spectral unmixing of the image, the result is overlaid with an intensity image.
(A) (B) (C) (D)
480 540 600 660 720
0
1
2
3
4
5
6
7
Fluorescein
Autofluorescence
Wavelength(nm) Fl. Intensity (A.U)
58
4.1.3. Fluorescence lifetime imaging
We developed a mosaic fluorescence lifetime imaging system that provides a large field
of view, using a femtosecond (fs) pulsed laser, for multi-mode optical imaging of small
animals. In order to evaluate the feasibility of this method, a model system study was
performed. Additionally, mosaic fluorescence lifetime imaging of the same mouse which
was used for a spectral imaging evaluation has been performed.
4.1.3.1. Ex vivo test of mosaic fluorescence lifetime imaging system
For the evaluation of the mosaic fluorescence lifetime imaging system, after the letters
“MISTI” were written on a white paper with 1mM fluorescein solution diluted by PBS
(pH 7.4), the region of interest on the paper was scanned by our mosaic fluorescence
lifetime imaging system (2cm x 3cm). fs pulsed laser light at 437.5nm with a beam
diameter of 1.8cm, generated through the same setup shown in Figure 3-8, was used for
the excitation of the fluorescein on the paper. Also, a 500nm longpass filter rejects the
excitation light from emission light in front of the time-gated intensifier CCD camera. A
total of six fluorescence lifetime images were obtained by mosaic fluorescence lifetime
imaging. And then, the images were merged into a single image for large field of view,
and 2x2 Gaussian filter was applied into the merged image in order to reduce the high
frequency at the boundary of the tiled images using ImageJ. Figure 4-4 shows the tiled
images and the merged image with a size of 2.7x3.8cm. In the figure, the lifetimes of
fluorescein are from 3.5ns to 4.5ns.
59
Figure 4-4. Large area image obtained by mosaic fluorescence lifetime imaging: (A)
Fluorescence lifetime images obtained for spatially combining image. After the letter was written
on a white paper with 1mM fluorescein solution (pH 7.4), the area (2x3cm) of interest in the
paper was scanned by mosaic fluorescence lifetime imaging with a step size of 1cm. (B) Image
reconstructed from the six images obtained by mosaic FLIM.
4.1.3.2. In vivo test of mosaic fluorescence lifetime imaging with a nude mouse
Additionally, mosaic fluorescence lifetime imaging of the same mouse which was used
for a spectral imaging evaluation has been performed. Figure 4-5 shows the fluorescence
lifetime image displayed in pseudocolor.
While the fluorescence lifetimes from a mouse skin display around 3~3.8ns, the
fluorescence lifetimes of fluorescein injected into tumor regions show approximately
3.8~4.5ns. In the image, the tumor regions with fluorescein are clearly distinguished from
other regions by the fluorescence lifetime difference. Also, the fluorescence lifetime
values of fluorescein are in good agreement with values in a literature (Gratton et al.,
2003).
(B)
(A)
1 Scan direction 6
1.8cm White paper
3.8cm
2.7cm
Lifetime(ns)
5.0
4.0
3.0
2.0
1.0
0.0
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Figure 4-5. Fluorescence lifetime image of a nude mouse with an implanted breast tumor: For
excitation of the injected dye (fluorescein), a 437.5nm fs pulsed laser with a 3.5cm beam diameter
(3.5cm), generated by the second harmonic of 875nm, was used. The mouse was scanned by the
system step by step for a large field of view. Two fluorescence lifetime images were acquired
sequentially. The images were merged into a single image. For each fluorescence lifetime image,
a total of 39 images have been acquired within 0 to 7800ps with a time step of 200ps.
4.1.4. Intra-vital confocal imaging
Before intra-vital confocal imaging, the optical fiber bundle probe is calibrated with a
calibration tool kits (Mauna Kea Tech.). And then, after a optical fiber bundle probe is
positioned inside a small animal by the holder, the intra-vital confocal images are
obtained. Figure 3-9 shows the schematic of the intra-vital confocal imaging mode.
The feasibility of intravital confocal imaging of a rat spine has been evaluated. After
euthanizing a rat, the rat spine was exposed. Then, autofluorescence images of the rat
spine were recorded for 8min by using an MKT S-probe (0 μm working distance). The
typical frame rate is 12 frames/sec. Figure 4-6 shows the experimental setup and results
obtained by using it. The images in Figure 4-6(B) clearly show convolutions of
neuromuscular junction (the first image at the first row), skeletal muscle fibers (the first
and second image at the second row), and vessels (the second and third image at the
second row).
Lifetime(ns)
0
4.0
2.0
61
Figure 4-6. Intra-vital confocal imaging of the rat spine: (A) Experiment setup. The arrow
indicates the optical fiber bundle probe which is positioned by the holder. (B) Intra-vital confocal
images of the rat spine. Several positions have been investigated during image acquisition.
4.1.5. Bioluminescence imaging
Using bioluminescence mode in this system, we imaged the presence of ATP and
enzymatic activity in nude mice. After luciferin was injected into the abdominal cavity of
the nude mice, we took a photographic image and then recorded the bioluminescence
signal for 2 minutes. Figure 4-7 shows the overlay image of the photographic and
bioluminescence image. These results demonstrate the capability to record
bioluminescence signals with good signal-to-noise and versatility, opening up the
possibility for further studies using this modality.
Figure 4-7. Bioluminescence image from engineered mice (normalized by highest value).
(A)
(B)
1
0
62
4.1.6. Scanning/wide-field two-photon excited fluorescence imaging mode
Scanning/wide-field two-photon excited fluorescence imaging mode was developed in
order to examine the details of molecules or tissues at deeper locations in multimode
optical imaging in vivo, outside the system. In this section, the evaluation of scanning
two-photon excited fluorescence imaging is performed. In addition, wide-field two-photon
excitation for multimode optical imaging in vivo is described in detail.
4.1.6.1. Scanning two-photon excited fluorescence imaging of a mouse liver
Two-photon excited fluorescence image of a mouse liver stained nonspecifically with
fluorescein is obtained by using our scanning two-photon excited fluorescence imaging
mode. In Figure 4-8, liver cells stained with fluorescein are shown. Femtosecond pulsed
laser light at 780nm was here used for the two-photon excitation of the fluorescein
(500μMol), and a shorpass filter (<700nm) and a bandpass filter (540±40nm) were
utilized for the emission light selection.
Figure 4-8. Two-photon excited fluorescence image of a mouse liver stained with fluorescein
(Nikon 20x, NA: 0.50).
100μm
63
4.1.6.2. Wide-field two-photon excitation in multimode optical imaging in vivo using
non-scanning imaging devices
In this section, the optical characteristics of wide-field two-photon excited fluorescence
imaging and contrast enhancement by wide-field two-photon excitation compared to one-photon
excited fluorescence imaging are demonstrated. In addition, the usefulness of
wide-field two-photon excited fluorescence imaging and the feasibility of multimode
optical imaging with two-photon excitation for small animals in vivo are examined. In the
investigation of the optical characteristics of wide-field two-photon excitation, power,
wavelength, numerical aperture, and depth dependence of wide-field two-photon
excitation with gallium corrole solution, sectioned mouse intestine, and a tissue phantom
are examined. Also, in vitro measurement of wide-field two-photon excited fluorescence
imaging is performed on an eye specimen taken from an alzheimer’s-disease mouse for
comparison between one-photon and wide-field two-photon excited fluorescence imaging.
Here, a flat-field correction method is applied in order to reduce artifacts due to Gaussian
beam profile of excitation laser light. In addition, vascularization around implanted tumor
regions of a nude mouse is examined by using wide-field two-photon and one-photon
excited fluorescence imaging, and then we compare these results in order to evaluate the
usefulness of wide-field two-photon excited fluorescence imaging for small animals in
vivo. Finally, we examine the feasibility of combination of wide-field two-photon
excitation and a multimode instrument using non-scanning in small animal imaging in
vivo.
64
4.1.6.2.1. Optical characteristics of wide-field two-photon excited fluorescence imaging
The dependence of fluorescence intensity on laser power has been examined with 50μM
sulfonated gallium corrole solution in order to evaluate the wide-field two-photon
excitation. A total of six wide-field two-photon excited fluorescence images were
acquired with the power range (200 ~ 400mW) of excitation laser at 795nm with a step
size of 50mW, and then fluorescence intensity of each image has been measured. After
that, curve fitting to the measured intensity along as a function of excitation power has
been performed. Figure 4-9(A) shows the measured intensity and fitted curve
(I=0.01xP2.01). In the figure, the fluorescence intensity increases nonlinearly with the
excitation power, suggesting a nonlinear process of fluorescence generation.
We have examined wavelength dependence of wide-field two-photon excited
fluorescence imaging with the mouse intestine sample. A wavelength region (780 to 910
nm) from the pulsed laser light was used for two-photon excitation of the sample.
Fluorescence intensity of the images obtained by wide-field two-photon excited
fluorescence imaging was measured at each wavelength. Figure 4-9(B) shows
fluorescence intensity versus wavelengths. In the figure, fluorescence intensity has a peak
at 790nm (Dickinson et al., 2003). In addition, photobleaching has been examined before
and after wide-field two-photon excited fluorescence imaging to confirm the accuracy of
the result. As a result, 0.03% photobleaching occurred during the experiment [intensity
(Arbitrary Unit): from 1541 to 1497].
65
Figure 4-9. Optical characteristics of wide-field two-photon excited fluorescence imaging: (A)
Power dependence of wide-field two-photon excited fluorescence of sulfonated gallium corroles.
(B) Wide-field two-photon excited fluorescence versus wavelength. We have examined the
wavelength dependence of wide-field two-photon excited fluorescence imaging of the mouse
intestine stained with Alexafluor 568 phalloidin. The excitation power at each wavelength was
250mW. The exposure time was 568ms. A total of 14 images have been acquired at different
wavelengths (780-910nm, step: 10nm). Here, an emission filter (620±60nm) and a 40x air
objective were utilized. (C) Wide-field two-photon excitation versus a numerical aperture. Wide-field
two-photon excited fluorescence has been investigated with a different numerical aperture of
objectives. Here, a fs pulsed laser source (780nm, 200mW) was used for the excitation of the
specimen, and 10x (NA: 0.45), 40x (NA:0.75), 40x (NA:1.30), and 60x (NA:1.45) objective were
used for collection of fluorescence. The rectangle represents the size of the field of view, 256μm
x 244μm, 64μm x 61μm, and 43μm x 41μm, respectively.
Characteristics of wide-field two-photon excitation due to numerical aperture were
investigated with the mouse intestine specimen. Figure 4-9(C) shows that the fitting
curve clearly shows nonlinearity (4th order) of the wide-field two-photon excitation to a
numerical aperture. The field of view does not depend significantly on numerical aperture
but is changed in proportion to the magnification of the objectives. Here, the fluorescence
intensity obtained with the 40x (NA: 1.30) objective is four times higher than that with
the 40x (NA: 0.75) objective.
A tissue phantom was constructed as a mimic for comparison of contrast and confocal
characteristics of one-photon, scanning two-photon, and wide-field two-photon excited
fluorescence imaging (Flock et al., 1992). The phantom is composed of a nude mouse
skin, a mixed gel (125μM fluorescein, 5% intralipid, and 1% microspere), and paraffin.
I =1.602NA4
100μm
(A) (B) (C)
0.4 0.6 0.8 1.0 1.2 1.4 1.6
0
10
20
30
40
50
60
70
80
FL. Intensity (A.U)
Numerical Aperture
780 800 820 840 860 880 900 920
200
300
400
500
600
700
800
Wavelength (nm)
Fl. Intensity (A.U)
200 250 300 350 400
500
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1500
2000
Power(mW)
Fl. Intensity (A.U)
I = 0.01× P2.01
66
After a thin sample of nude mouse skin is put on a slide glass in a chamber, the mixed gel
is loaded on the mouse skin. Finally, paraffin was filled in the chamber in order to
prevent evaporation from the skin [Figure 4-10(D)].
With the constructed tissue phantom, a total of 52 images within 0~153μm with a step
size of 3μm have been obtained by one-photon, scanning two-photon, and wide-field
two-photon excited fluorescence imaging respectively.
Figure 4-10. Comparisons of depth dependence in one-photon, scanning two-photon, and wide-field
two-photon excited fluorescence imaging of a tissue phantom: (A) Scanning two-photon
excited fluorescence image: fs pulsed laser light source at 780nm was used for excitation. The
average laser power is 50mW. (B) Wide-field two-photon excited fluorescence image. Excitation
wavelength was 780nm. The average laser power was 200mW. (C) One-photon excited
fluorescence image: Excitation at 488±30nm and power of 140μW. (D) Constructed phantom. (E)
Intensity profiles over the selected regions along a depth.
Figure 4-10(A), (B), and (C) shows the scanning two-photon, widefield two-photon, and
one-photon excited fluorescence image at the same focal plane, respectively. In the figure,
-20 0 20 40 60 80 100 120 140 160
0.0
0.2
0.4
0.6
0.8
1.0
Scanning two-photon
Wide-field two-photon
One-photon
depth (um)
Fl. Intensity (A.U)
(A) (B) (C)
Mouse skin
0 depth mark Glass slide
paraffin
125μM Fluorescein+
5% intralipid+1%microsperes+agarose gel
(D) (E)
50μm 50μm 50μm
67
the intensity profiles of the microsphere selected by a solid rectangle have been examined
along a depth for contrast comparison. Figure 4-10(E) shows the profiles along the depth
(z-axis). According to the profiles, a full width at half maximum of the intensity profile in
the scanning two-photon excited fluorescence imaging is narrowest (approximately
10μm), while the full width at half maximum in wide-field two-photon excited
fluorescence imaging shows approximately 60μm, which is significantly less than that
(over 160μm) in one-photon excited fluorescence imaging.
4.1.6.2.2. In vitro Measurements
Figure 4-11(A) and (B) shows the results obtained by one-photon and wide-field two
photon excited fluorescence imaging of a mouse intestine section stained with Alexafluor
568 phalloidin.
Figure 4-11 Wide-field two-photon excited fluorescence images of a mouse intestine: (A) One-photon
excited fluorescence image of the flamentous actin of the mouse intestine stained with
Alexafluor 568 phalloidin with Texas red filter set (Ex: 560nm±40nm, Em: 630nm±60nm) and
838ms exposure time. (B) A wide-field two-photon exited fluorescence image (up and right) of
the regions selected by the white dotted rectangle; the image (bottom and right) has been
corrected by a Gaussian mask (down and left). The power of the excitation laser light at 800nm
was 300mW with an exposure time of 458ms. (C) Profile of cross-sections of the original and the
corrected image along a line (A’).
(A) (B) (C)
50μm
50μm
Original
Corrected
Pixel
F.L intensity
0 50 100 150 200 250
1.0
1.5
2.0
2.5
3.0
3.5
4.0
68
In the wide-field two-photon excited fluorescence images, filamentous actin is clearly
shown as in one-photon excited fluorescence image. In addition, in order to reduce
artifacts by the Gaussian profile of the excitation beam, the image is corrected by the
mask generated by the previously described method. In Figure 4-11(B), the mask, the
original, and the corrected image are displayed together. The profiles across the original
and the corrected image are shown in Figure 4-11(C). The profile of the wide-field two-photon
excited fluorescence image corrected by the constructed mask shows uniform
baseline compared to that of the uncorrected image. Wide-field two-photon excited
fluorescence imaging of an eye specimen of an alzheimer’s mouse was also performed.
Figure 4-12. One-photon and wide-field two-photon excited fluorescence imaging of an eye
specimen of an alzheimer’s mouse: (A) One photon excited fluorescence image (ex: 560nm, em:
630nm). (B) Wide-field two-photon excited fluorescence image (ex: 830nm, em: 630nm, power:
300mW) and the Gaussian corrected image (white rectangle).
Figure 4-12 shows one- and two-photon excited fluorescence image of the eye specimen
at the same depth. First of all, plaques on the eye were found using one-photon excited
fluorescence imaging, and then wide-field two-photon excited fluorescence image of the
same region was acquired. In addition, flat-field correction of the wide-field two-photon
excited fluorescence image was performed in order to enhance contrast. Figure 4-12(B)
shows the wide-field two-photon excited fluorescence image. Here, more plaques in the
50μm
(A) (B)
50μm
69
wide-field two-photon excited fluorescence image are clearly observed than in one-photon
excited fluorescence image.
4.1.6.2.3. Multimode optical imaging with wide-field two-photon excited fluorescence
imaging in vivo
Wide-field two-photon excited fluorescence imaging of a small animal in vivo has been
performed. After 100μM gallium corrole solution was injected into tumor regions of a
euthanized mouse [Figure 4-13(A)], the regions were investigated using one-photon
excited fluorescence imaging. Thin vessels are observed in one-photon excited
fluorescence image as shown in Figure 4-13(B). In addition, the thin vessels have been
examined by wide-field two-photon excited fluorescence imaging, and then the image
was flat-field corrected [Figure 4-13(C)].
In the Figure 4-13(C), the thin vessels are more clearly shown than in the one-photon
excited fluorescence image [Figure 4-13(B)]. Finally, after the injection of microsphere
into the subcutaneous regions indicated by the green dotted circle, a total of 40 one-photon
and wide-field two-photon excited fluorescence images with different depths were
obtained.
Figure 4-13(D) and 4-13(E) show images at four focal planes (10μm, 10.5μm, 13.5μm,
and 17μm) beneath the skin. In these figures, while several microspheres are only seen in
wide-field two-photon excited fluorescence images at 13.5μm and 17μm [Figure 4-13(E)],
they are not easily seen in the one-photon excited fluorescence images at those focal
planes [Figure 4-13(D)].
70
Figure 4-13. Wide-field two-photon excited fluorescence imaging and multimode optical imaging
with wide-field two-photon excitation of a nude mouse with implanted tumors after subcutaneous
injection of corroles and miscrosphere solution: (A) Euthanized nude mouse. Corroles and
miscrosphere solution were injected into breast tumor regions selected by a red solid circle and
the subcutaneous of regions selected by a dotted green circle, respectively. (B) One-photon
excited fluorescence image of the red circled regions. Light (wavelength: 425nm ± 50nm)
filtered from a Hg lamp have been used for one-photon excitation. The bandpass of the emission
filter was 620nm± 40nm. (C) Wide-field two-photon excited fluorescence image corrected by a
Gaussian mask. 800nm fs laser light (250mW) was been utilized for two-photon excitation (Em:
620nm± 40nm). The field of view is approximately 75μm. (D) One-photon excited fluorescence
images at different depths. 10μm, 10.5μm, 13.5μm, and 17μm beneath the skin around the regions
selected by the green dotted circle (E) Two-photon excited fluorescence image at different
depths: 10μm 10.5μm, 13.5μm, and 17μm. (F) Spectral and fluorescence lifetime images with
wide-field two-photon excitation. A total of 13 spectral images within a spectral range from
500nm to 610nm, with 10nm bandwidth and 10nm step size, were obtained. The images were
analyzed using our developed software. Spectral signatures of fluorescence from microspheres
and autofluorescence are shown in the graph at the bottom and left side. Also, a total of 22 images
within 0~4.2ns with step size 0.2ns, have been obtained. The fluorescence lifetime image has
been constructed using the first order exponential decay fitting method. The graph at the bottom
and right side shows the time-decay curves of fluorescence and autofluorescence within the
selected regions.
(F)
(A) (B) (C)
(D) (E)
Intensity (A.U) Wavelength (nm) Time (ns)
Intensity (A.U)
0
1 1
0
500 560 620 0 2 4
Lifetime(ns) 0
3
2
1
Microsphere
Autofluorescence
Microsphere
Autofluorescence
71
In addition, spectral and fluorescence lifetime imaging of the subcutaneous regions with
wide-field two-photon excitation have been performed in order to separate the
microspheres from autofluorescence. In Figure 4-13(F), the pseudo-color mapped images
represent a spectral classification (top-left) and a fluorescence lifetime image (top-right),
respectively. For the spectral classification, we selected two spectral signatures shown in
the figure (bottom-left). In the spectral signatures, while the spectral signature of
fluorescence from the microspheres has a peak at 560nm, the autofluorescence has a
broad spectral signature. Also, Figure 4-13(F) (upper-left) shows the classified image
based on the selected signatures. Here, a red color represents fluorescence of
microspheres and a green color represents autofluorescence. In the fluorescence lifetime
image, while fluorescence lifetime of microspheres is approximately 1.8~2.3ns, the
fluorescence lifetime of autofluorescence is below 1.2ns. The graph (bottom-right) in
Figure 4-13(F) shows that the intensity of fluorescence from microspheres and
autofluorescence decays over a period of 0~4.2ns. The exponential decay coefficients of
the fluorescence and autofluorescence display 2.15ns and 0.78ns.
4.1.6.2.4. Partial Summary
In this section, we have shown the optical characteristics of wide-field two-photon
excitation, usefulness of wide-field two-photon excitation in small animal imaging in vivo,
and feasibility of non-scanning multimode instruments with the wide-field two-photon
excitation. This excitation method can provide better contrast and more penetration depth
than one-photon excited fluorescence imaging, and it is more resistant to artifacts due to
animal movement than scanning two-photon excited fluorescence imaging, although the
72
confocal capability and contrast in wide-field two-photon excited fluorescence imaging is
less than those in scanning two-photon excited fluorescence imaging.
Figure 4-9(A), (B), and (C) show the optical characteristics of wide-field two-photon
excitation. The plot of the fluorescence intensity versus wide-field two-photon excitation
power increased with a slope of 2.01. This is in good agreement with typical power
dependence of two-photon excitation. In addition, the excitation spectra of wide-field
two-photon excitation of Alexafluor 568 are similar to those in the literature (Dickinson
et al., 2003). These results demonstrate that wide-field two-photon excitation of
fluorescence behaves like that using laser light focused to a diffraction-limited spot.
Also, in the tissue phantom study for comparison of images acquired by wide-field two-photon,
one-photon, and scanning two-photon excited fluorescence imaging,