by Samira Karimelahi - University of Toronto T-Space...Samira Karimelahi Masters of Applied Sciences Graduate Department of Electrical and Computer Engineering University of Toronto

Combinational multiphoton scanning microscopy andmultiphoton surgery of mouse arteries

by

Samira Karimelahi

A thesis submitted in conformity with the requirementsfor the degree of Masters of Applied Sciences

Graduate Department of Electrical and Computer EngineeringUniversity of Toronto

Copyright c© 2011 by Samira Karimelahi

Abstract

Combinational multiphoton scanning microscopy and multiphoton surgery of mouse

arteries

Samira Karimelahi

Masters of Applied Sciences

Graduate Department of Electrical and Computer Engineering

University of Toronto

2011

Preliminary investigations were carried out in order to explore the potential of laser-

stimulated capillary growth in a blood vessel-on-a-chip. To fulfill the project objective,

a series of experiments in both directions of two photon fluorescence imaging and laser-

semitransparent materials interaction were performed. A purpose-built two-photon flu-

orescence imaging resolution was tested by imaging 1 µm diameter fluorescent beads.

Also, the potential of fluorescence imaging in the waveguide writing field as well as the

biological field was studied. Further, for laser ablation on the mouse artery loaded in the

microfluidic channel, the processing window was found such that the damage induced by

femtosecond laser just affects the artery, not the other interfaces of the microfluidic chip.

At the end, the result of laser trepanning on the mouse artery wall combined with two

photon fluorescence imaging was shown. These results will be useful for more advanced

biological study such as angiogenesis.

ii

Acknowledgements

I would like to express my deep gratitude to my supervisor, Professor Peter Herman, for

his guidance and encouragement. In addition, I would like to thank Dr. Jianzhao Li

for his help and for his training. Thanks to our collaborators, Professor Axel Guenther

and Professor Steffen Sebastian Bolz. Special thanks also go to my family: my parents

Najmeh Rafiepour Nabiollah Karimelahi, for their patience and support. They always

believed in me and helped me not to feel lonely even though I was far from them. I

would also like to thank all the people in photonics group at University of Toronto for

their assistance and friendship these last two years, especially Nima Zareian and Ladan

Abolghasemi.

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Contents

1 Introduction 1

1.1 Thesis objectives . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4

1.2 Chapter by chapter outline . . . . . . . . . . . . . . . . . . . . . . . . . . 5

2 Background 7

2.1 Laser interaction with transparent materials . . . . . . . . . . . . . . . . 7

2.1.1 Nonlinear ionization . . . . . . . . . . . . . . . . . . . . . . . . . 8

2.1.2 Femtosecond laser modification inside transparent materials . . . 10

2.2 Cell and tissue disruption by femtosecond laser . . . . . . . . . . . . . . . 12

2.2.1 Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 21

2.3 Two-photon fluorescence imaging . . . . . . . . . . . . . . . . . . . . . . 24

2.3.1 Mechanism of multiphoton fluorescence microscopy . . . . . . . . 24

2.3.2 Architecture of two-photon fluorescence microscope . . . . . . . . 27

2.3.3 Multiphoton imaging applications . . . . . . . . . . . . . . . . . . 30

3 Experiment 36

3.1 Femtosecond laser system . . . . . . . . . . . . . . . . . . . . . . . . . . 37

3.2 Beam delivery system . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 38

3.3 Purpose-built two-photon fluorescence imaging setup . . . . . . . . . . . 41

3.3.1 Hardware of the fluorescence microscope system . . . . . . . . . . 41

3.3.2 TCSPC laser microscope software . . . . . . . . . . . . . . . . . . 49

iv

3.4 Sample preparation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 52

3.4.1 Mouse blood vessel loaded in the microfluidic chip . . . . . . . . . 52

3.4.2 Fluorescent microspheres . . . . . . . . . . . . . . . . . . . . . . . 57

4 Results and discussion 59

4.1 Two photon fluorescence imaging of microspheres . . . . . . . . . . . . . 60

4.1.1 TPI by two dry lenses with different numerical apertures . . . . . 61

4.1.2 TPI by oil-immersion lens . . . . . . . . . . . . . . . . . . . . . . 67

4.2 Two photon fluorescence imaging of the waveguides inside the fused silica 73

4.3 Two photon fluorescence imaging of the optical fiber . . . . . . . . . . . 77

4.4 Breakdown threshold at various microfluidic chip interfaces . . . . . . . . 79

4.4.1 Damage threshold: glass-air interface . . . . . . . . . . . . . . . . 81

4.4.2 Damage threshold: glass-PDMS interface . . . . . . . . . . . . . . 86

4.4.3 Damage threshold: glass-MOPS interface . . . . . . . . . . . . . . 89

4.4.4 Comparison between damage threshold results for different interfaces 91

4.4.5 Bubble formation threshold inside MOPS solution . . . . . . . . . 95

4.5 Micromachining on the mouse artery wall . . . . . . . . . . . . . . . . . . 99

4.5.1 Combination of trepanning and two photon fluorescence imaging . 105

5 Conclusion 109

5.1 Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 109

5.2 Future directions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 110

A Light propagation in the matter 112

B Two photon absorption probability 115

Bibliography 116

v

List of Acronyms

3D: Three dimensional

AOM: Acousto-optic modulator

CCD: Charge coupled device

FWHM: Full width at half maximum

IR: Infrared

LBO: Lithium triborate

MOPS: 3-(n-Morpholino)Propanesulfonic Acid

NA: Numerical aperture

PMT: Photomultiplier tube

TTL: Transistor-transistor logic

vi

List of Tables

3.1 Optical focusing parameters related to three lenses. . . . . . . . . . . . . 44

4.1 Comparison of calculated beam spot size and depth of focus as well as the

measured lateral and axial length of the fluorescent beads. . . . . . . . . 70

4.2 Comparison of calculated lateral and axial resolution according to the

Rayleigh Resolution criteria definition. . . . . . . . . . . . . . . . . . . . 73

4.3 Laser exposure parameters for writing waveguides inside fused silica. . . . 74

4.4 Threshold pulse energy and irradiance for different microfluidic chip inter-

faces in single and a multiple-pulse laser exposure. . . . . . . . . . . . . . 93

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List of Figures

2.1 Schematic models of different kinds of the photoionization according to

Keldysh parameter, γ [31]. . . . . . . . . . . . . . . . . . . . . . . . . . . 9

2.2 Timescale of different physical phenomena happening during a femtosec-

ond laser interaction with matter. Although the pulse duration is fem-

tosecond, the permanent changes in the matter occur in a microsecond

scale. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11

2.3 Array of laser static exposure showing index modification produced by

(a) different laser energies and various number of pulses, 1.4 NA, and 110

fs pulse duration at 1 kHz in Corning 0211 glass [31].(b) Sodalime glass

irradiated with 60 fs pulse duration laser and 5.5 nJ pulse energy with

different repetition rates and number of bursts per spot using oil immersion

objective with NA=1.4 [36].(c) Inclination of the cover slip makes 10 nm

difference in the focus point for two adjacent hole created by 527 nm laser

beam focused through the 1.3 NA objective lens in Corning 0211 [37]. . . 13

2.4 Transmission of 800 nm and 110 fs laser pulses focused by 0.65 NA lens

inside fused silica as a function of the laser energy [31]. . . . . . . . . . . 14

2.5 Diagram of important parameters in laser-tissue interaction [40]. . . . . . 15

2.6 Absorption spectra for the three dominant components: water, hemoglobin

and melanin in tissue [41]. . . . . . . . . . . . . . . . . . . . . . . . . . . 16

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2.7 Time resolved laser ablation photos of corneal tissue taken at 313 ns, 4.5

µs, 29 µs, and 25 ms after laser irradiation [42]. . . . . . . . . . . . . . . 17

2.8 (a) Ablation lines with five different pulse energies on fluorescently-labeled

actin fibers. (b) Fluorescence intensity profile with respect to position

along sample [50]. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 21

2.9 Results of the laser machining in fixed rat neocortical tissue (a) Array of

static exposure with varying pulse energy and number of pulses . (b) Cross

section image of the volume removed as a result of the single shot with

0.65 µJ pulse energy. (c) Repeated line cuts with 0.1 mm/s scan speed

and 0.5 µJ pulse energy. (d) Side view of the fixed cortical tissue that

shows double cut. The first cut removed an area of 1 mm2 with depth

of 200 µm, and the second cut removed an area of 0.25 mm2 with a final

depth of 360 µm [51]. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22

2.10 Two-photon fluorescence images of the brain’s blood vessel disruption us-

ing femtosecond laser (a) with high energy that leads to hemorrhage, (b)

with lower energy that generates extravasation, (c) and with several num-

ber of pulses that leads to cloting [52]. . . . . . . . . . . . . . . . . . . . 23

2.11 Comparison between the volume of excitation in (a) single photon ex-

citation and (b) two-photon excitation. In single photon excitation the

fluorescence signal can be seen from the whole path of the laser beam in

(a), but in two-photon excitation shown in (b) the fluorescence signal is

coming from the much smaller focal volume [12]. . . . . . . . . . . . . . . 25

2.12 (a) Wavelength distribution of fluorescence and second harmonic signal

(b) isotropic emission of the fluorescence signal [12]. . . . . . . . . . . . . 26

2.13 Dependence of the excitation process on axial distance for one photon and

two-photon excitation [11]. . . . . . . . . . . . . . . . . . . . . . . . . . . 28

2.14 Two-photon fluorescence imaging set up [12]. . . . . . . . . . . . . . . . . 29

ix

2.15 2PF/SHG microscopy of the (a) untreated, and (b) controlled rat’s artery

rings. In (b) lindane usage caused morphological change in the artery wall.

(c) and (d) are the zoomed-in view of (a) and (b), respectively [73]. . . . 31

2.16 Multiphoton imaging of the vascular distribution of (a) normal and (b)

tumor blood vessels [74]. . . . . . . . . . . . . . . . . . . . . . . . . . . . 32

2.17 Temporal evolution of drug delivery technique imaged by two-photon flu-

orescence imaging [75]. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 33

2.18 Combination of two-photon microscopy and femtosecond laser microsurgery

on a breast carcinoma cells single layer. (a) Two-photon image of a sin-

gle layer of live breast carcinoma cells before irradiation with a laser. (b)

Two-photon image right after irradiation with a single pulse at 280 nJ

pulse energy causing fluorescence signal lost in the targeted cell. Scale

bars are 20 µm [77]. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 34

3.1 Fiber chirped-pulse amplification arrangement of the fiber fs laser [2]. . . 38

3.2 Beam delivery setup for the femtosecond fiber laser. TM is a turning

mirror and FM is a flipping mirror. See text for detailed explanation of

components. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 39

3.3 Burst of the laser pulses generated with the high repetition rate MHz

ultrashort laser system using AOM (a) to produce 100% on/off laser beam

modulation and (b) to create an envelop of 80% duty cycle which is shown

in dotted square wave [84]. . . . . . . . . . . . . . . . . . . . . . . . . . . 40

3.4 Npaq Control Assembly showing various jumpers. The jumper JP1 was

changed from the default position of 1-2 to 2-3. . . . . . . . . . . . . . . 41

3.5 Depth correction of refraction of the light according to the Snell’s law at

the interface of two materials with different refractive indices n1 and n2. . 43

3.6 Two-photon fluorescence imaging setup. See text for detailed explanations

of components. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 45

x

3.7 Hardware block diagram of the imaging system (modified from [86]). . . . 46

3.8 Principle of the TCSPC measurement [88]. . . . . . . . . . . . . . . . . . 48

3.9 Three trigger pulses that determine pixel, line, and frame [86]. . . . . . . 48

3.10 Control and Analysis tab showing different features of the TCSPC Laser

Microscope Software [87]. . . . . . . . . . . . . . . . . . . . . . . . . . . . 50

3.11 (a) 3D display of the sample two-photon fluorescence image. (b) 3D in-

tensity profile [87]. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 51

3.12 Image acquisition software view showing a 2D image of multiple 1 µm

microspheres. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 53

3.13 Schematic illustration of the different steps of fabricating PDMS stamp.

Modified from [93]. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 55

3.14 Microfluidic chip structure and position of the blood vessel. A 1-2 mm

length blood vessel is loaded to an artery inspection area via an artery

loading well using suction pressure [94]. . . . . . . . . . . . . . . . . . . . 56

3.15 Images of 1 µm diameter fluorescent beads (a) under SEM, and (b) under

a bright field microscope with 100× objective lens. . . . . . . . . . . . . 58

4.1 Cross section TPI of the 1 µm fluorescent beads recorded with 40X-0.65

NA. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 62

4.2 The axial intensity profile of one of a fluorescent bead represent in (a)

ImageJ and in (b) MATLAB (red line). . . . . . . . . . . . . . . . . . . . 63

4.3 Scanning laser beam through the microsphere beads mounted on the mi-

croscope slide. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 64

4.4 The x-y image of the fluorescent beads recorded with a 0.65 NA lens: each

pixel is equivalent to 0.5 µm size. . . . . . . . . . . . . . . . . . . . . . . 65

4.5 The x-y image of the fluorescent beads recorded with a 0.65 NA lens, each

pixel is equivalent to 0.25 µm size. Scales are similar for both directions. 66

xi

4.6 The lateral intensity profile of a fluorescence sphere (1 µm diameter) ob-

served and represented by a Gaussian curve (red line) with MATLAB tools. 67

4.7 An x-y image of the fluorescent beads by a 100X 0.9 NA objective lens

where each pixel is equivalent to 0.25 µm × 0.25 µm area. Scales are

similar for both directions . . . . . . . . . . . . . . . . . . . . . . . . . . 68

4.8 An x-z image of the fluorescent beads by a 0.9 NA objective lens, each

pixel is equivalent to 0.5 µm × 0.5 µm area. Scales are the same for both

directions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 69

4.9 An x-y image of the 1 µm diameter fluorescent beads recorded with the

1.25 NA oil-immersion lens for a pixel size equivalent to 0.5 µm. . . . . . 71

4.10 The cross-section images (x-z) of the 1 µm diameter fluorescent beads with

the 1.25-NA oil-immersion lens where each pixel size is equivalent to (a)

0.5µm and (b) 0.25 µm. . . . . . . . . . . . . . . . . . . . . . . . . . . . 72

4.11 Optical microscopic image of waveguides written inside fused silica (a),

transverse (xy) TP images of waveguides, and Cross sectional (xz) TP

images of the waveguides. . . . . . . . . . . . . . . . . . . . . . . . . . . 75

4.12 Waveguide cross-sectional phase contrast microscopic images for circular,

parallel and perpendicular polarizations laser beam at 1 MHz repetition

rate, 175 nJ pulse energy, and 0.75 mm/s scan speed [2]. . . . . . . . . . 76

4.13 Single mode fiber optic taped on a microscope slide for TP laser image. . 77

4.14 TPI of a optical fiber with a 40X-0.65 NA dry lens (a) cross section (xz)

and (b) top view (xy). . . . . . . . . . . . . . . . . . . . . . . . . . . . . 78

4.15 TPI of a fiber soaked in oil with a 100X-1.25 NA oil immersion lens a)

cross section (xz) and b) top view (xy). . . . . . . . . . . . . . . . . . . . 79

4.16 Optical microscopic images of femtosecond laser line scan on the top sur-

face of the cover slip interface with air. Each average power line scan is

repeated for scan speeds from 0.2 to 50 mm/s. . . . . . . . . . . . . . . . 82

xii

4.17 Microscopic image showing array of femtosecond laser static exposure.

Each spot on the figure is corresponding to a specific average power and

number of laser pulses. The static exposure was sensitive to even 1 µm

displacement in the focusing position. The lowest threshold is taken to

be the damage threshold. The exposure points are separated by 20 µm in

each direction. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 83

4.18 Exposure zones repeated for static exposure in different focal positions

offsets of -2,-1, 0, 1, 2 µm on the cover slip to find the lowest damage

threshold. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 84

4.19 Minimum number of pulses corresponding to the specific pulse energy re-

quired to induce damage in a cover slip top surface with a 1045 nm 300 fs

at 1 MHz laser beam. The solid line is a guided to the average of four sets

of data. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 85

4.20 Optical microscopic image showing array of femtosecond laser exposure

on a glass bottom surface. Each exposure point is separated by 20 µm in

each direction. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 86

4.21 Minimum number of pulses required to induce damage on a cover slip

bottom surface with a 1045 nm 300 fs at 1 MHz laser beam. The solid line

is a guided to the average of four sets of data. . . . . . . . . . . . . . . . 87

4.22 Comparison between laser damage threshold on the bottom and top sur-

faces of the cover slip for glass-to-air interfaces. . . . . . . . . . . . . . . 88

4.23 Considering the boundary conditions for the collimated light, the exiting

surface has lower breakdown threshold than the entering surface because

of the stronger electric field at the exiting surface [95]. . . . . . . . . . . 88

4.24 An array of laser exposures on the cover slip-PDMS interface with varying

number of pulses and average powerfor 1045 nm 1 MHz 300 fs laser radiation. 89

xiii

4.25 Average of minimum number of pulses required to induce damage in the

cover slip-PDMS interface in each power for 1045 nm 1 MHz 300 fs laser

radiation. Error bars indicate the standard deviation. . . . . . . . . . . . 90

4.26 An array of laser exposures on the cover slip-MOPS interface with varied

number of pulses and average power for 1045 nm 1 MHz 300 fs laser radiation. 91

4.27 Average of minimum number of pulses required to induce damage in the

cover slip-PDMS interface in each power for 1045 nm 1 MHz 300 fs laser

radiation. Error bars indicate the standard deviation. . . . . . . . . . . . 92

4.28 Comparison between damage threshold of different interfaces at microflu-

idic chip for various exposure condition. . . . . . . . . . . . . . . . . . . . 94

4.29 Free electron density versus normalized irradiance with respect to the

threshold irradiance. This plot is provided for 100 fs pulse duration at

three different laser wavelength [8]. . . . . . . . . . . . . . . . . . . . . . 96

4.30 The minimum number of laser pulses inducing bubble formation inside the

MOPS solution versus average power. The solid line is a guide. . . . . . . 97

4.31 Comparison of all Breakdown thresholds at different exposure conditions. 98

4.32 Mouse artery wall loaded in the microfluidic chip. The laser beam ablated

the blood vessel in the x direction. . . . . . . . . . . . . . . . . . . . . . 100

4.33 The results of the scanning laser beam scanning across artery wall. Images

at different focus positions are shown here for each set of laser parameters

because of the irregular shape of artery wall in the present experiments.

Window show the laser exposure conditions. . . . . . . . . . . . . . . . . 101

4.34 Scanning laser beam in the z direction (vertical to blood vessel surface)

on the artery wall. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 102

4.35 30 µm hole created by laser power of (a) 80 mW laser beam and (b) 70 mW.103

xiv

4.36 Selected video frames with a red number on top observing the laser-tissue

interaction while laser trepanning with 70 mW laser power. The bubble

formation were observed in a few frames i.e. frame number 6. . . . . . . 104

4.37 Cross sectional two photon fluorescence image of the unstained artery wall.

Scale is the same for both directions. . . . . . . . . . . . . . . . . . . . . 105

4.38 256×256 pixel image of the blood vessel (every pixel is 0.5µm) with the

exposure parameters of 1.13 nJ pulse energy at 1 MHz repetition rate a)

Top view (x-y) view of the blood vessel which is dyed with Propidium Io-

dide. b) Cross section (x-z) image of the blood vessel with both Propidium

Iodide (right image) and Fura-Red stained (left image). . . . . . . . . . . 106

4.39 The CCD camera images of the artery wall (a) before and (b) after laser

exposure. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 107

4.40 Two photon fluorescence imaging of the artery wall after laser trepanning

exposure. The diameter of the hole is around 80 µm. . . . . . . . . . . . 108

xv

Chapter 1

Introduction

Femtosecond lasers make it possible to drive the nonlinear processes inside materials. The

pulses with short duration and high intensity can induce material structural modifications

in both absorptive and transparent materials. In these regimes, materials will behave

nonlinearly that makes the multiphoton absorption possible [1–4].

The first femtosecond laser machining was demonstrated in 1994. One aspect that

tried to be improved is to increase the resolution of the laser interaction. Focusing

ultrashort pulses under high numerical aperture lenses makes nanometer scale ablation

possible [1].

Transparent material modifications by femtosecond lasers have drawn significant at-

tention in recent years [5]. High intensity short pulses can induce localized modification

and damage inside the material via multiphoton ionization, tunneling, and avalanche

ionization. As the nonlinear processes cause breakdown in the material, the damage is

confined to the focal volume. In the other words, only in the sub diffraction-limited

focus diameter, the laser intensity is high enough to induce nonlinear excitation. The

laser modification in the material can be moved to write three dimensional structures

inside glass as for applications such as direct writing of the optical waveguides [6] and

three-dimensional binary data storage [7]. In addition to formation of photonics devices,

1

Chapter 1. Introduction 2

femtosecond lasers have been widely used in biological areas. In biological manipulation,

it is important to have localize effects in order to minimize collateral damage in the sam-

ple. Femtosecond lasers are able to create this confined interaction in biological systems

with minimum collateral damage [8, 9]. Consequently, ultrashort lasers can be used for

nanosurgery and to study biological dynamics by targeting selective parts of the cell,

multiple cells, or tissue [8, 10].

Multiphoton absorption induced by femtosecond lasers also have applications in flu-

orescence imaging. Ultrashort laser pulses in the mid-infrared spectrum can excite the

bio-material with photon energy equivalent to that in the UV range. Because of the non-

linear excitation only in the focal volume, three dimensional image sectioning is obtain-

able without any pinholes or other spatial filters [11]. Multiphoton fluorescence imaging

that is based on the nonlinear excitation have advantages such as deeper penetration

depth, lower photo damage and higher photo bleaching threshold, over confocal imaging

which is based on the linear absorption [12]. These properties make nonlinear fluorescence

3D imaging a good substitute for confocal microscopy in areas such as neuroscience [13].

Multiphoton imaging has been used in a variety of imaging tasks. In biological ap-

plications using combination of laser imaging with machining on biological structures,

this method will result in a better understanding of cellular responses to an external

disruption [8, 10]. Two photon fluorescence imaging has been applied in wide range of

research like the study of embryonic development [14], intracellular free calcium activ-

ity [15], neuronal plasticity [16], and angiogenesis [17]. Multiphoton fluorescence imaging

can also be a useful tool to analyze and visualize waveguides in optical circuits.

One of the challenges in working on biological samples during femtosecond laser ex-

posure is how to hold the sample and keep the sample alive. During laser exposure,

it is important to fix the sample and oxygenate it. One way to address these needs is

Microfluidic devices. Microfluidic devices are useful in handling small sample sizes and

integrating multiple processes required for lab-on-a-chip (LOAC) experiments. These


properties make microfluidic chips appropriate for analyzing single cells, cellular struc-

tures, and tissue [18]. The samples like C.elegan or blood vessels can be immobilized

inside the channel in the microfluidic. In addition, biological solutions like MOPS can

keep the environment appropriate to keep the sample alive. Also, as there is just a trans-

parent cover slip on top of the sample, this way of holding the sample will not impede

the imaging process.

The combination of femtosecond laser surgery with two photon fluorescence imaging

on a sample which is trapped in a microfluidic chip offers an interesting new research

direction. Microfluidic chip technology in combination with diversified femtosecond lasers

interaction physics offer interesting investigations in worm biology [19].

Another interesting areas to explore is to take advantage of this combination and

study other physiological process like angiogenesis, which is the growth of the new blood

vessels from existing vessels. The blood vessel can be loaded inside a microfluidic device

and stimulated by a femtosecond laser while taking the two photon fluorescence imaging.

The aim of this thesis is to explore the laser-blood vessel interaction by exposing

tissue samples with a femtosecond laser and combining micromachining with two photon

fluorescence imaging. This work will open the door to more investigations in the novel

study of angiogenesis, in which the laser surgery should be performed on the mouse artery

wall while it is alive.

In order to investigate our imaging set up characterization, a part of this thesis is

dedicated to the two photon fluorescence imaging of microspheres with 1 µm diameter.

Taking the two photon fluorescence image of the fluorescent beads with different objec-

tive lenses will help to find the appropriate objective lens for each application to offer

high resolution imaging. Further, two photon fluorescence imaging is applied to image

the waveguides inside fused silica to show the potential of the fluorescence imaging in

waveguide characterization and analysis.

This work was completed together with Dr. Jianzhao Li provided training on working


with the laser and two photon fluorescence imaging. This work was also a team project

with the medical group of Professor Steffen Sebastian Bolz of Physiology department for

vessel study and Professor Axel Guenther of the Mechanical Engineering for microfluidic

chip design and fabrication.

1.1 Thesis objectives

In this thesis, we take advantage of both lab-on-a-chip devices fabricated by Professor

Axel Guenther’s group and the powerful femtosecond laser available in our lab. The

objective is to demonstrate our two photon fluorescence imaging capabilities, both in

imaging the waveguides and the mouse artery loaded into a microfluidic chip and studying

the femtosecond laser interaction with the mouse artery wall.

Experiments that were done to fulfill the project objective were as followings:

1. Study of two photon fluorescence resolution by taking images of 1 µm fluorescently

dyed spheres with three different objective lenses.

2. Record the fluorescence images of single mode optical fibers and waveguides written

inside the fused silica.

3. In order to study the femtosecond laser interaction with the mouse artery wall, the

following steps were completed:

• Determine the damage threshold of the microfluidic chip components at differ-

ent interfaces as well as the bubble formation threshold in the MOPS (physi-

ologic salt solution).

• Micromachining on the blood vessel wall to demonstrate controllable damage

induced by the femtosecond laser.

• Two photon fluorescence imaging of both unstained and stained blood vessels.


• Combination of the laser trepanning on the mouse artery wall and two photon

fluorescence imaging.

1.2 Chapter by chapter outline

The overview of each chapter is presented by the following outline:

• Chapter 2 “Background” reviews the light-transparent materials interaction focus-

ing on the nonlinear processes induced by the femtosecond laser. The ultrashort

pulses can be used to modify sample structures in both semiconductors and bio-

logical tissues. Also, examples of femtosecond laser applications in cell and tissue

disruption are given. Further, two photon fluorescence imaging mechanisms and

architectures as well as its applications are presented in this chapter.

• Chapter 3 “Experiment” presents the femtosecond laser system and the beam deliv-

ery path. Also, our purpose-built fluorescence imaging setup is explained in detail

for both hardware and software aspects. Moreover, the sample preparation meth-

ods including both mouse artery loaded in the microfluidic chip and fluorescent

microspheres are given.

• Chapter 4 “Results and discussion” reviews varies experiments. First, results of

the microsphere two photon fluorescence imaging by three different objective lenses

are reviewed and compared. Then the two photon fluorescence imaging of the

waveguides inside fused silica and single mode optical fiber are given. Also, the

breakdown threshold measurement results for different interfaces of the microfluidic

chip are compared and presented. Further, micromachining on the mouse artery

loaded in the microfluidic chip and its combination with two photon fluorescence

imaging are reviewed.


• Chapter 5 “Conclusion” presents a summary of this research work and its signifi-

cance as well as possible future directions.

Chapter 2

Background

Ultrashort pulse duration lasers have unusual properties which make them useful tools

for science and applications [20]. Such laser pulses are very short in time for probing

fast physical and chemical processes. Their wide spectral bandwidth can be useful for

dense wavelength division multiplexing (DWDM) in optical networks [21] and selective

excitation of the fluorescent dyes in multiphoton fluorescence imaging [22]. Ultrahigh

peak intensity created in femtosecond pulse duration can drive multiphoton absorption

and nonlinear interaction with materials that makes the ultrashort laser useful in a wide

range of applications like material ablation [23], material structuring inside glass [24–26],

two-photon imaging [16], and nanosurgery [8, 27,28].

2.1 Laser interaction with transparent materials

The advent of high-power pulsed lasers makes it possible to study the laser-induced

breakdown inside transparent materials [29].

Femtosecond pulses in comparison with nanosecond and picosecond lasers with the

same average power have higher intensity and can provide electric field that exceeds

the electric field that holds electrons in the valence band. Therefore, the electron can

be excited and brought from the ground state to the excited state. In this regime of

7

Chapter 2. Background 8

intensity, interactions between the laser and the material is nonlinear. In other words,

material which is transparent to the laser in low intensity will become opaque with high

intensity laser light.

As found in more detail in Appendix A, when the laser intensity is low, the material

polarization is a linear function of the electric field while at high intensity this relation

becomes nonlinear and the refractive index will be a function of the laser intensity.

Dispersion, diffraction, and aberration are examples of linear effects, and self focusing

and plasma defocusing are as a result of the nonlinear phenomena [2].

Photoionization and avalanche ionization are two different classes of nonlinear ab-

sorption that can take place in the material interaction with high intensity lasers [29,30].

This nonlinear absorption of the laser energy can result in permanent damage in the

material. The advantage of the ultrashort lasers is that the damage induced inside the

transparent materials is localized, because only in the focal volume will the intensity be

high enough to cause damage. One can use these structural changes, like refractive index

modification, to write small structures in order to create integrated optical components

inside the transparent materials [29].

2.1.1 Nonlinear ionization

Ultrashort laser pulses with high intensities can deposit energy to the matter via various

nonlinear excitation mechanisms. The electron can be promoted from the valence band to

the conduction band as a result of the photoionization and the avalanche ionization [31,

32]. Incident beam energy is transferred to the matter, first as electrons are ionized and

then transfer their high energy to the lattice via this collision with the ions.

For transparent materials, a single photon of visible or infrared light does not have

enough energy to excite an electron, so multiphotons are required to promote the electron.

Electrons can be directly excited via photoionization depending on the laser frequency

and intensity, following either of two different paths: multiphoton ionization and tunnel-


ing ionization. The value for the Keldysh parameter, γ, which is given by Eq. (2.1), will

determine which one of these two processes will take place:

γ =ω

e

√mecn�0Eg

I, (2.1)

where ω is the laser frequency, me is the effective electron mass, I is the laser intensity

at the focal point, c is the speed of light, e is the fundamental electron charge, n is the

linear refractive index, �0 is the permittivity of free space, and Eg is the bandgap energy.

According to Keldysh, photoionization will be multiphoton when γ > 1.5 and will be

tunneling when γ < 1.5, and will be via combination of these two processes when γ is

about 1.5 (Fig. 2.1). The photoionization rate depends on the laser intensity [31].

Figure 2.1: Schematic models of different kinds of the photoionization according to

Keldysh parameter, γ [31].

Another class of nonlinear absorption is avalanche ionization where an electron in

the conduction band can be promoted to a higher level by absorbing several photons

sequentially. When the energy of the electron exceeds the band gap energy plus the

conduction band minimum energy, the electron can excite another electron in the valence

band collisionally. These two electrons can then excite other electrons after they are

accelerated by the strong electric field of the laser to high kinetic energy. The rate of

the growth of the electron density, N, in the conduction band as a result of the impact

ionization is according to:


∂N

∂t= ηN, (2.2)

where η is the avalanche ionization rate.

There should be an excited electron in the conduction band to begin the avalanche

ionization. These initial electrons called seed electrons can be provided via thermal

excitation, ionized impurity, multiphoton or tunneling ionization [31,33]. Nonlinear ion-

ization can create the high density electron plasma that can strongly absorb laser energy.

Because of the typical spatial Gaussian shape of the laser intensity, the density of elec-

trons is high in the center and low in the wings of the beam. As electron density has

an inverse relation with the refractive index, this plasma can defocus the beam as it

propagates in the matter [31,33].

2.1.2 Femtosecond laser modification inside transparent mate-

rials

Laser energy deposited inside transparent materials via nonlinear absorption can be high

enough to cause permanent damage and material modification.

The physics of the femtosecond laser interaction with the matter is simpler than with

picosecond pulses because the time that an electron absorbs a photon is much shorter

than the time scale needed for transferring energy from electron to the lattice. In other

words, in the femtosecond regime the laser beam energy will heat the electron distribution

before being transferred to the lattice via electron-phonon scattering. As long as the laser

pulse is entering to matter, the number of the electrons in the conduction band is going

to increase. When the density of the electrons reaches the critical plasma density, plasma

will absorb most of the light, while at higher plasma density, the plasma region is going

to reflect most of the light [31, 34]. Only after the laser pulse has passed, energy will

be transferred to the lattice to cause the localized permanent change in the structure or


even create a void [31,35].

According to Fig. 2.2, electrons transfer energy to the lattice is typically on the time

scale of picoseconds. In a couple of nanoseconds, shock waves separate from the hot focal

point, and in the microsecond scale, heat will diffuse out of the focus point [8].

Figure 2.2: Timescale of different physical phenomena happening during a femtosecond

laser interaction with matter. Although the pulse duration is femtosecond, the permanent

changes in the matter occur in a microsecond scale.

The probability of absorbing light is expected to be proportional to IN in the material

with the band gap (Eg) equivalent to N photons energy satisfying Nhν = Es. But

experiments show that the threshold intensity does not depend on the material band

gap. So, femtosecond lasers can be used in machining a wide range of the materials.

Pulse duration, focusing numerical aperture, and the repetition rate are three param-

eters that affect the damage threshold intensity [36].

One of the complications in measuring damage threshold inside the transparent ma-

terial is self focusing. As a result of the self focusing, spatial and temporal properties

of the beam are changing inside the material. The self focusing threshold is a function

of the peak power and not the intensity. As power increases, the self focusing will in-

crease, until it reaches the critical power (Eq. (2.3)) in which self focusing balances the


diffraction and creates a filament. The critical power is given by:

Pcr =3.77λ2

8πn0n2, (2.3)

where λ is a laser wavelength, and n0 and n2 are linear and nonlinear refractive index of

the materials respectively. In order to avoid self focusing, one can use a high NA lens and

low power to get high intensity at the focal volume. On the other hand, one consequence

of a high focus lens is aberration which makes it difficult to reduce the spot size. It

has been shown that for materials with refractive indices between 1.3 and 2, NA=0.65

will narrow rays to less than 100 nm, which is smaller than the diffraction limited spot

size [29].

There are different methods like optical microscopy or transmission to measure the

material damage threshold. One way is to form an array of static laser exposures cre-

ated by varying laser parameters such as number of pulses and laser power is shown in

Fig. 2.3 [31,36,37]. Using optical microscopy, one can observe the refractive index change

in the material. In the transmission method, the laser power passing through the sample

is measured as power is changed gradually from low to high. At the power high enough

to induce damage inside the sample there will be reduced transmission power due to the

absorption of the laser energy by material structural modifications (Fig. 2.4) [29].

2.2 Cell and tissue disruption by femtosecond laser

Soon after the first demonstration of the Ruby laser in 1960, biomedical uses of lasers

started and developed towards wavelengths covering a wide range from UV (shorter than

visible wavelength) to IR (longer than visible wavelength) [38].

Laser-tissue interaction widely depends on the irradiance parameters and tissue prop-

erties like absorption and scattering coefficient, heat capacity, and thermal conductivity.

Laser parameters like energy pulse, repetition rate, wavelength, and beam spot size will


Figure 2.3: Array of laser static exposure showing index modification produced by (a)

different laser energies and various number of pulses, 1.4 NA, and 110 fs pulse duration at

1 kHz in Corning 0211 glass [31].(b) Sodalime glass irradiated with 60 fs pulse duration

laser and 5.5 nJ pulse energy with different repetition rates and number of bursts per spot

using oil immersion objective with NA=1.4 [36].(c) Inclination of the cover slip makes

10 nm difference in the focus point for two adjacent hole created by 527 nm laser beam

focused through the 1.3 NA objective lens in Corning 0211 [37].


Figure 2.4: Transmission of 800 nm and 110 fs laser pulses focused by 0.65 NA lens inside

fused silica as a function of the laser energy [31].

affect the reaction of the tissue to the incident beam [39]. The chart in Fig. 2.5 presents

important parameters in laser-tissue interaction.

The reflection, refraction, scattering, and absorption properties of tissue change with

the incident light wavelength. Although tissue has a complex structure and varied

chemical composition, it can be modeled by its dominant components such as water,

hemoglobin and melanin. According to absorption coefficients of these materials (Fig.

2.6), for the incident wavelength between 0.6 and 1.2 µm, tissue is considered transparent.

Researchers have tried to understand the dynamics of the laser-tissue interaction. For

example in [42], the temporal evolution of the ablation crater in corneal tissue have been

obtained. A 30 ps pulse duration and 1.1 mJ pulse energy coming from the mode-locked

Nd:YLF laser oscillator after 120 roundtrips amplification was used irradiate corneal

tissue. The first snapshot shown in Fig. 2.7, is taken 313 ns after corneal tissue had

been exposed. The shock front can be seen because of the induced refractive index


Figure 2.5: Diagram of important parameters in laser-tissue interaction [40].


Figure 2.6: Absorption spectra for the three dominant components: water, hemoglobin

and melanin in tissue [41].


change. Expansion of the vaporized gas and annular deformation of the tissue surface is

observable after 4.5 and 29 µs. The last exposure taken at 25 ms shows the remaining

ablation crater.

Figure 2.7: Time resolved laser ablation photos of corneal tissue taken at 313 ns, 4.5 µs,

29 µs, and 25 ms after laser irradiation [42].

Among lasers with different pulse duration, femtosecond lasers have been the most

interested when applied for biomedical imaging. Femtosecond lasers have become one of

the most useful tools for precise microsurgery due to the low energy threshold of bubble

formation. Bubble creation in the cell results in stretching and rupturing of the cell

membrane that finally will kill the cell. In order to observe the photodisruption in real


time, one can take advantage of time resolved two-photon spectroscopy [8, 22,43].

The high peak intensities and short time duration of the pulses lead to efficient and

rapid ionization of tissue before energy can be lost. Photoionization can be induced

by multiphoton and/or tunneling pathways depending on the laser frequency, duration,

and intensity. These processes will create quasi free electrons in the conduction band

which will generate plasma. The plasma formation in tissue, called laser induced optical

breakdown, plays a significant role in plasma ablation and photodisruption [8]. Ultra-

short pulses at the focal plane can exceed the electric field binding valence electrons, and

as a result of this optical breakdown, micro-plasma will be created at the focal plane.

The created plasma will absorb further energy from the laser pulse to cause strong tem-

perature and pressure gradients at the focal volume. Secondary effects arising from the

plasma formation is shock-wave and cavitation bubble creation. For the appropriate laser

source parameter this laser-tissue interaction will result in the precise tissue cutting with

controllable damage [22,43,44].

Increasing the pulse energy can increase the ablation efficiency, but after some thresh-

old, because of the exponential growth of plasma density, the ablation efficiency will fall

off. The plasma at the surface of the tissue will act as a shell that will absorb and scatter

the laser radiation and shield against deeper laser penetration [45].

One way of having a localized laser interaction inside of biological structures is to

tightly focus ultrashort laser pulses by means of a high numerical aperture lens. The

nonlinear absorption of high peak intensity in the femtoliter focal volume will confine the

damage to that small volume [8]. In order to drive laser ablation and modification on

biological samples, one can use femtosecond lasers with high repetition rate on the order

of a few MHz with energy levels just above the energy level require for nonlinear imaging

and well below the optical breakdown threshold. Another way is to use ultrashort lasers

with low repetition rate like 1 kHz and with pulse energies slightly above the ablation

threshold. In the first approach, thermal accumulation effects arise that induce cell lysis


and ablation of tissue, while in the second approach, each high pulse energy will induce

damage [8, 44].

Femtosecond lasers have the advantage of low optical breakdown threshold in trans-

parent materials that makes fs lasers the tool of choice for precise machining and surgery.

At typical intensities, light passes through transparent material without interaction or

ionization. For higher intensities, because of the high photon flux density, the interaction

probability of several photons simultaneously with the same molecule increases. There-

fore, multiphoton absorption causes ionization of the transparent material and creates

seed electrons for avalanche ionization which results in high density electron-ion plasma.

Laser pulse energy is stored in the plasma as free negative and positive charges with their

kinetic energy in the order of tens of picoseconds, electrons and ions recombine which

result in large amount of the energy being released that will result in breaking the tensile

force of the material around the focus position. Depending on the pulse energy, a non-

equilibrium thermal condition can lead to microexplosion and shockwave. Cavitation

bubbles created as a mechanical side effect further drive complex dynamics [8, 46,47].

Femtosecond laser localized machining happens when the electron density is below

the critical value which is calculated to be 1021cm−3 [48]. So, it is necessary to under-

stand laser interaction with the low-density plasma. Chemical changes, thermomechnical

processes, and heating are consequences of laser interaction with low-density plasma [8].

Two factors that minimize the laser affected zone should be taken into consideration.

First, the focusing condition, and second, the pulse energy. The laser beam should

be focused through a high NA lens to produce sufficient intensity to induce nonlinear

absorption just at the focal volume. As total energy deposited to the matter determines

the strength of the side effects like shockwaves, it is better to keep the pulse energy

low [8]. Moreover, when a train of the pulses hit the target in the time scale shorter than

the time needed for heat to diffuse out of the focal volume, the heat accumulation effect

will happen. This cumulative effect can lead to significant thermal and chemical effects.


Although chemical effects from single femtosecond pulses with low energy for the cells

can be negligible, repetitive pulses can result in useful or harmful chemical reactions [49].

In oder to create controllable damage on biological system by femtosecond lasers, it

is important to do a systematic damage threshold experiment. The damage threshold

may vary for each sample as the optical properties of each biological system is different.

Damage threshold measurements can be done on the subcellular, cellular, and tissue level

and can be carried on in the different form of the laser machining as will be illustrated

below.

The relation between femtosecond laser pulse energy and the subcellular dissection has

been studied by Heisterkamp et al. [50]. They applied a train of 100 fs laser pulses with

nanojoule range pulse energy from a titanium-sapphire laser system at the repetition of

1 kHz. Pulses were focused on the sample with a 1.4 NA oil immersion lens. The results

in Fig. 2.8 demonstrate five ablation lines with various pulse energies were detected from

the fluorescence image of the actin network of a fixed endothelial cell. The scan speed

for each line was 0.7 µms

[50].

Another point of interest is to measure the laser-tissue damage threshold by histology.

Examples of forming holes, surface channels, and deep tissue removal in brain tissue are

shown in Fig. 2.9. Femtosecond lasers can be used to cut and image the brain tissue at

the same time. Changing the exposure parameters will result in three different regimes

of ablation which was demonstrated by Tsai et al. [51]. These three different regimes are:

the static ablation by varied pulse energy and number of exposure pulses while scanning

the sample to show the relation between pulse energy and spatial extend of the ablation,

line cutting that was done in the fixed cerebellar tissue, and millimeter scale slab cut as

demonstrated in Fig. 2.9.


Figure 2.8: (a) Ablation lines with five different pulse energies on fluorescently-labeled

actin fibers. (b) Fluorescence intensity profile with respect to position along sample [50].

2.2.1 Applications

One application of femtosecond lasers is to simulate human disease in animals like rodents

and mice to create a model for research. For example, Nishimura et al. [52] have

used ultrashort lasers to create novel models of neurovascular disease such as strokes

in the mouse brain that rely on controllable laser damage and without disturbing the

surrounding tissue area. To demonstrate this goal, 100-fs laser pulses were focused on

the lumen of blood vessel within the 500 µm of the cortex [52].

Depending on the laser energy deposited to the blood vessel, three classes of distur-

bances can happen as is described in Fig. 2.10 [52]. One disturbance is blood plasma


Figure 2.9: Results of the laser machining in fixed rat neocortical tissue (a) Array of

static exposure with varying pulse energy and number of pulses . (b) Cross section image

of the volume removed as a result of the single shot with 0.65 µJ pulse energy. (c)

Repeated line cuts with 0.1 mm/s scan speed and 0.5 µJ pulse energy. (d) Side view of

the fixed cortical tissue that shows double cut. The first cut removed an area of 1 mm2

with depth of 200 µm, and the second cut removed an area of 0.25 mm2 with a final

depth of 360 µm [51].


extravasation which is toxic for the neurons. Second, ischemia happens as a result of the

stop in the blood flow. Third, hemorrhages occur because of the vessel rupture. Such

surgical disruption can illustrate wide ranging physical conditions from changing blood

flow to neural death.

Figure 2.10: Two-photon fluorescence images of the brain’s blood vessel disruption us-

ing femtosecond laser (a) with high energy that leads to hemorrhage, (b) with lower

energy that generates extravasation, (c) and with several number of pulses that leads to

cloting [52].

Another application of the lasers is in surgery. Lasers have been widely used for eye

surgery for decades [53]. Human eyes are transparent in the visible and near IR range and

they are easily accessible for surgery. For eye correction, femtosecond laser can be applied

to the eye to cut a portion of the cornea and reshape it to fix its focus position [54].

Moreover, femtosecond lasers have application in study of biological systems as they

have the ability to dissection the subcellular scale. So, femtosecond lasers can be used

to selectively disrupt, for example, part of the neuronal circuit to study its neuronal


behavior such as studying axonal regrowth [55]. Another example of the femtosecond

laser application in this area is to study the structure of the mitochondria [56].

2.3 Two-photon fluorescence imaging

Two-photon excitation (TPE) processes were first proposed by Goppert-Mayer in 1931

who won the Nobel Prize in physics for working on nuclear shell physics [11]. She theo-

retically showed that multiple photon absorption can cause excitation which normally is

induced by a single photon [12].

The first demonstration of multiphoton microscopy (MPM) was by the Watt W.

Web group a decade ago. MPM is based on the excitation of fluorescence within the

small volume inside the sample. Although the primary signal source for MPM is two-

photon excited fluorescence, one can do imaging based on the second harmonic (SHG)

and third harmonic generation (THG). Another form of nonlinear imaging is anti-Stokes

Raman scattering (CARS) which requires two synchronized laser sources at different

wavelengths [12].

2.3.1 Mechanism of multiphoton fluorescence microscopy

Fluorescence microscopy can be based on linear or nonlinear excitation. In one photon

absorption, the incident frequency should be the same as the resonance frequency of the

molecule. This raises the electron from the ground state to excited state, from which

it relaxes to the electronic ground state and emits a lower energy photon [12]. Two-

photon fluorescence refers to the excitation of a fluorophore when two-photons arrive

within a time window of an attosecond. These two photons cooperatively provide the

energy needed to excite fluorescence. As a result, excitation can take place in the infrared

spectral range [13,57].

Nonlinear optical microscopy is more capable than confocal microscopy in biological


imaging. In the confocal microscope, the light source should be in the near-UV range in

order to excite single-photon electronic transitions in various fluorophores. On the other

hand, nonlinear optics can take advantage of the infrared and near infrared light that

can penetrate more deeply into the tissue even up to 500 µm, while the 2 photon energy

matches to excite the same fluorescent states [11, 58]. Moreover, multiphoton imaging

has the advantage that photobleaching is confined to a very small focal volume of about

few femtoliters ( Fig. 2.11). As the incident wavelength is directly proportional to the

spatial resolution, the confocal microscopy has better resolution over the multiphoton

microscopy. [12, 59].

Figure 2.11: Comparison between the volume of excitation in (a) single photon excitation

and (b) two-photon excitation. In single photon excitation the fluorescence signal can be

seen from the whole path of the laser beam in (a), but in two-photon excitation shown

in (b) the fluorescence signal is coming from the much smaller focal volume [12].

In the fluorescence microscopy, the sample is illuminated with the light source the

wavelength that can excite fluorophore inside a specimen. The emission spectra will be

collected with the appropriate detector. This method is useful for the three dimensional

study of biological systems and their dynamic properties.

As just few of the biological structures have primary fluorescence, it is necessary to


attach fluorescent dye (fluorophore) in order to be able to probe the fluorescence signal.

The component of a molecule which causes a molecule to absorb energy of a particular

wavelength and emit energy at a different wavelength is a fluorophore. There are several

properties like low photobleaching, large absorption cross section, and low phototoxcity

to cells that define an ideal fluorophore [12].

The emission wavelength of the fluorophore is usually less than the incident wave-

length and higher than one half of the wavelength (Fig. 2.12a). Fluorescence signals

are emitted in all directions around the laser interaction volume (Fig. 2.12b), but sec-

ond or third harmonic signals are directional as they should meet the phase matching

condition [12].

Figure 2.12: (a) Wavelength distribution of fluorescence and second harmonic signal (b)

isotropic emission of the fluorescence signal [12].

If the frequency of the incident light is one half of the atom resonance frequency,

and the photon flux density is high enough, then two photons can be absorbed by the

same fluorophore simultaneously and induce an excitation process. Also, the laser pulse

duration should be shorter than the atom relaxation time which is of a time scale of

10−9 s. In this case, when an atom absorbs one photon, it does not have enough time

to relax, so it can absorb another photon. As a result, picosecond and femtosecond near

infrared lasers are appropriate light source for two-photon microscopy [11,22,60].


As it is shown in detail in Appendix B, one important factor in two-photon absorption

probability, na, is the two-photon cross section which is different for each fluorophore. A

larger two-photon cross section, σ2, results in higher two-photon absorption rate, so it

is important to select the fluorophore with large σ2 that can be excited by the available

laser source [61].

Multiphoton absorption depends nonlinearly on the intensity. This intensity depen-

dence makes multiphoton absorption localized. As shown in Eq. (B.4), two-photon ab-

sorption is proportional to the square of the intensity. This quadratic dependence orig-

inates from the need for two photons to absorb and induce an excitation. On the other

hand, for one photon excitation, the linear relation between intensity and absorption

typically cause non-localize excitation (Fig. 2.13) [11]. This nonlinear dependence will

permit optical sectioning in two-photon imaging. So, by scanning the beam inside the

sample one can build the three dimensional image without any need to pinhole. An ap-

propriate detector can collect the fluorescence signal coming from the interaction volume

in the sample [11].

2.3.2 Architecture of two-photon fluorescence microscope

An arrangement for two-photon fluorescence imaging is shown in Fig. 2.14. One impor-

tant component for two-photon fluorescence imaging is the laser. The choice of the laser

source is critical because the appropriate one focused with the high NA lens should have

the high photon flux density to increase the probability of absorbing two photons in the

small time window. Although it is possible to induce two-photon absorption even with

continuous laser, femtosecond and picosecond lasers are the appropriate laser sources for

imaging. For a short pulsed laser, low power will offer high intensity in a tight focus and

yet remain less harmful for the cell and tissue as the net energy deposited to the sample

is proportional to the average power. One of the most common lasers for two-photon

microscopy is the titanium-sapphire laser systems that provide high repetition rate and


Figure 2.13: Dependence of the excitation process on axial distance for one photon and

two-photon excitation [11].


femtosecond pulse duration with moderate average power. Both picosecond and fem-

tosecond laser sources can be used to trigger two-photon absorption, but for getting the

same level of the fluorescence signal, picosecond laser should have higher average power

which this cause photodamage in the sample [11].

Electronics to control and synchronize beam scanners and detectors are crucial el-

ements in microscopy which determines the speed of the capturing frame. Computer-

controlled motion stage or galvanometric scanning mirrors are common scanner device.

High numerical aperture objective lenses are also necessary components in order to focus

tightly and get high intensity [11, 12]. Finally, appropriate detectors that have the high

efficiency to collect the fluorescence signal are required. The detector selection parame-

ters are spectral range, electronic noise level, cost, readout speed, and quantum efficiency.

Photomultiplier tubes (PMT), avalanche photodiode (APD), and charge-coupled detec-

tors are three main detectors using in fluorescence microscopy [62]. The laser scanning

Figure 2.14: Two-photon fluorescence imaging set up [12].

confocal microscope set up is similar to the two photon microscope, but the laser source

is different. Moreover, for confocal microscope it is necessary to have the pinhole in


front of the detector to achieve optical sectioning for constructing three dimensional im-

ages. It is common that people buy the commercial confocal microscope and modify it

to get multiphoton fluorescence imaging. Also, scanning mirror should be modified to

one reflecting the new laser source (infrared) [11].

2.3.3 Multiphoton imaging applications

Application in biology

Multiphoton microscopy (MPM) has been used widely in biology to study physiology,

morphology, and cell-to-cell interaction. MPM is one of the powerful tools in biology to

image thick tissue even in the live animal, useful to monitor the dynamics of biochemical

processes [63]. Neuroscientist have applied MPM to monitor the calcium dynamic depth

in the brain tissue [12, 64–68] to study neuronal plasticity [16]. Also, study of the dy-

namics of calcium deep can be useful to study neurodegenerative disease models in both

brain slice [69] and in live mice [70–72].

Another example of MPM is to image the blood vessels to monitor the effect of

the lindane. Lindane was used as a disinfectant and insecticide in agriculture until the

mid-70s, but because of its toxicity it was banned. Using two-photon fluorescence in

combination with SH imaging can show the impact of this toxic material on the arterial

tissue. Fig. 2.15 shows a nonlinear image of the artery ring before and after it was treated

by lindane. The image of the treated rat’s artery shows the alternation in the artery wall

that becomes wavier as a result of lindane. This experiment shows that the waviness of

the laminae is increased by roughly 10 % in arteries of treated rats in comparison with

the control one [73].

Multiphoton fluorescence imaging can be also useful in the study of angiogenesis,

which is the growth of new blood vessel from an existing one, vessel remodeling and

vessel maturation. From multiphoton imaging, the differences between angiogenic blood


Figure 2.15: 2PF/SHG microscopy of the (a) untreated, and (b) controlled rat’s artery

rings. In (b) lindane usage caused morphological change in the artery wall. (c) and (d)

are the zoomed-in view of (a) and (b), respectively [73].


vessels and normal blood vessels are observable. Also, by means of this method, it will

be feasible to analyze changes in the blood vessel walls and to quantify the number and

spacing of the blood vessels as well as permit measurement of the vessel diameter and

length [17,74]. Further, the branching patterns are observable. As is shown in Fig. 2.16,

normal microvessels have well-organized architecture with dichotomous branching while

the tumor vessels are dilated, tortuous, saccular, and heterogeneous in their spatial dis-

tribution [74].

Figure 2.16: Multiphoton imaging of the vascular distribution of (a) normal and (b)

tumor blood vessels [74].

Another example of a MPM application is to monitor the temporal evolution of dis-

ruption in blood-brain barrier as a result of specific drug delivery method like ultrasound

enhanced with microbubble contrast agents. In order to permit visualization of the vas-

culature, mice were injected intravenously with fluorescent dyes. Fig. 2.17 shows the

real time two-photon fluorescence images of the vascular system. Each image is recorded

at various times of the treatment. Immediately after taking the first frame at t=0 s,

the drug delivery process started. Using MPM makes it possible to monitor the vessel

diameter during the drug delivery process [75].

Also, two-photon fluorescence imaging can be combined with femtosecond laser micro-

nanosurgery to make a powerful ”seek-and-treat” tool. In other words, MPM is acting as

an accurate non-invasive monitoring tool which can be helpful to visualize the region of


Figure 2.17: Temporal evolution of drug delivery technique imaged by two-photon fluo-

rescence imaging [75].

interest and shows the result of the precise femtosecond surgery [76, 77]. The combined

application of femtosecond lasers for both imaging and manipulation of biological samples

can be used for analysis and treatment of various diseases in addition to in vivo monitoring

of disease progression [52,78–80].

The combined imaging and microsurgery capabilities of femtosecond lasers illustrated

using breast carcinoma cells grown in a single cell layer as a sample with the fluorescent

cell viability dye labeling. In order to do that, the cell was imaged before and after laser

exposure. Cell images are shown in Fig. 2.18 [77].

Because of the localized damage of the femtosecond laser, it is possible to induce

photodamage in just one cell while adjacent cells are intact. The evidence of the photo

damage is the loss of the fluorescence signal observed in the image. A pulse energy

increased from 160nJ to 280 nJ will cause the fluorescence signal from targeted cell to be

lost. Because the size of the cell is much larger than the focal spot, it was claimed that

this signal drop is not due to photobleaching [77].


Figure 2.18: Combination of two-photon microscopy and femtosecond laser microsurgery

on a breast carcinoma cells single layer. (a) Two-photon image of a single layer of live

breast carcinoma cells before irradiation with a laser. (b) Two-photon image right after

irradiation with a single pulse at 280 nJ pulse energy causing fluorescence signal lost in

the targeted cell. Scale bars are 20 µm [77].

Application in imaging of the waveguides in the glass

In addition to the wide application of femtosecond lasers scanning microscopy in the biol-

ogy field, they can be applied to characterize, analyze, and visualize optical waveguides.

In this case, one can characterize waveguides using the same laser used for the fabrication

process.

One example of the femtosecond laser application in waveguide fabrication is to fab-

ricate a waveguide to bridge between two existence waveguides. So, one can find the

exact position of the two waveguides and then fabricate the bridge precisely in between.

Because both imaging and fabrication can be done with the same laser system, there is

no need to relocate the sample between two systems and therefore avoid realignment the

sample.

Also, combination of femtosecond laser microscopy and spectroscopy (microtroscopy)

diagnostics have potential applications in micro/nanofabrication. This will provide guid-


ance for in-situ laser trimming or post-processing as well as real-time feedback for con-

trolling laser fabrication process [81].

Chapter 3

Experiment

Experiments carried out for this work include two-photon fluorescence imaging and fem-

tosecond laser machining. Two-photon fluorescence imaging was performed on 1 µm

fluorescent beads and on waveguides written inside fused silica glass, single mode optical

fiber, and mouse blood vessel. Also, femtosecond laser micromachining was applied to

measure the damage threshold at different interfaces of the microfluidic chip and mouse

artery loaded into the microfluidic chip channel. To perform these experiments, fem-

tosecond laser pulses were guided to the sample via the beam delivery system. In order

to record two-photon fluorescence images of the sample, three types of equipment were

used: an optical setup to collect fluorescence, electronics hardware to count the number

of photons detected in time via the time correlated single photon counting technique and

software to control the stage and capture the image. The details of each part will be

explained later in this chapter.

Initially the femtosecond laser system used for both two-photon fluorescence imaging

and threshold measurement experiments is described. Then, the beam delivery system

for femtosecond laser micromachining will be explained. Moreover, details of the purpose-

built two-photon florescence imaging setup, both in hardware and software areas will be

given. Finally, the sample preparation including loading the mouse blood vessel in the

36

Chapter 3. Experiment 37

microfluidic chip and preparing microsphere fluorescence beads will be discussed.

3.1 Femtosecond laser system

The laser used for this thesis work is fiber-chirped pulse amplification (CPA)-fs laser

system (IMRA µJewel D-400-VR) which creates a pulse duration of around 300 fs. The

output repetition rate is variable between 100 kHz and 5 MHz and the average power is

500 mW. In this range of repetition rate, one can obtain maximum pulse energy varied

from 100 nJ to 5 µJ for repetition rate of 100 kHz to 5 MHz. This range of operation fills

the gap between high energy 1-250 kHz Ti:Sapphire regenerative amplifiers and low pulse

energy 80 MHz Ti:Sapphire oscillators [2]. The fiber chirped-pulse amplification technol-

ogy used in this laser is shown in Fig. 3.1. The seed pulses generated by Ytterbium-fiber

laser oscillator are expanded by using a fiber stretcher prior to entering the fiber amplifier

in order to avoid nonlinear damage. The amplified pulses will be then compressed by the

free space grating to achieve short pulses. The beam quality factor M2, which is defined

as the beam parameter product (product of the beam radius measured at the beam waist

and the beam divergence half-angle measured in the far field) divided by λ/π, can be

determined using the CCD camera [2,82]. In order to perform such a measurement, one

can focus the laser beam on the sample surface via a high NA lens, and find the position

where the spot size is minimized at the CCD image to measure the beam waist according

to the number of pixel occupied and pixel size. The next step will be to move the lens to

the position where the beam waist becomes larger by a factor of√

2; the axial distance

between these two points shows the Rayleigh range (zR). Subsequently, the beam diver-

gence half-angle can be calculated via√

λπzR

. For our system, M2 was calculated by my

colleagues to be 1.3.


Figure 3.1: Fiber chirped-pulse amplification arrangement of the fiber fs laser [2].

3.2 Beam delivery system

The beam delivery setup for our laser system is shown in Fig. 3.2. The stretched pulse

of a few hundred picoseconds are guided to the compressor via mirrors labeled TM1 and

TM2 which both have a high reflectivity at 1045 nm. The compressor box is outside the

laser head so that the prism inside the compressor is accessible for optimizing position

and correcting dispersion for every repetition rate. After the compressor, the 300 fs

pulses can be attenuated by a half wave plate and polarizer controlled by the computer

for exposure control. The polarization of the laser beam after the polarizer is horizontal

and parallel with the table surface. Depending on the structure to be written, the beam

can pass through the acousto optic modulator, AOM (Neos 23080-3-1.06-LTD), or just

skip that by flipping mirrors FM1 and FM2. In the AOM a piezoelectric transducer is

attached to a tellurium dioxide crystal. The transducer is vibrated by AC electric signal

causing acoustic waves to propagate through the tellurium dioxide crystal and generate

a periodic refractive index grating. This grating induces diffraction in the laser light

propagating through the crystal to generate a first order beam. The first order can take

0 to 60% of the incident power while the remaining power applied in the zero order beam

which is in the same direction as the incident beam. One can modulate the laser beam

power by turning the AC AOM signal on or off to control the existence of the first order

beam.

A second harmonic arrangement can be inserted to produce 522 nm wavelength light.


The laser beam is directed through the objective lens by mirrors TM6 and TM7.

The objective lens is mounted on the Z motion (Aerotech ALS130) stage and the

sample is on the XY motion stage (Aerotech ABL1000). The XYZ stage was controlled

by a computer via G-code software (Aerotech). One can use back illumination to record

the image of the sample with a FireWire-interfaced CCD camera (Sony XCD-X710)

which has a zoom lens (Computar L5Z6004). The light from a fiber bundle illumination

source can be directed onto the sample back by using a prism. The light collected with

the objective lens is directed to the CCD camera via mirrors FM3 and TM9. Also, as

the minimum beam spot size on the CCD camera corresponds to focusing on the sample

surface (aside from small correction due to laser beam divergence), one can use the CCD

camera to position the laser focus close to the sample surface.

Figure 3.2: Beam delivery setup for the femtosecond fiber laser. TM is a turning mirror

and FM is a flipping mirror. See text for detailed explanation of components.

In our group, the AOM has been widely used to modulate the pulse train and write

continuous arrays of refractive index voxels for writing Bragg Grating waveguides [83].

The AOM modulates the laser to create burst trains of pulses with controllable duty

cycle. In this case, the off time of the laser needs to be minimized so laser power is not

greatly reduced (Fig. 3.3). On the other hand, in the static exposure, the laser on-time

should be controllable on the order of a few µs to make it possible to expose samples


with a few pulses in one spot on the sample for sufficient refractive index change. In

order to get a few number of pulses in one position, the laser burst should be on in a

microsecond time scale which is not achievable for the default motion controller drive

(Aerotech A3200, Npaq) control board assembly. Consequently, the default settings of

the motion controller drive, which control the AOM via computer, needed to be modified.

Our motion controller drive is capable of controlling up to six axes of motion. The I/O

capabilities of the drive include a 16 channel opto-isolated digital I/O interface, four

16-bit analog inputs, two 16-bit analog outputs, and a single axis Position Synchronized

Output interface (PSO, or laser firing). The position synchronized output (PSO) was

used to control the AOM. The motion controller drive’s jumpers were set to the default

at the factory and could be changed to accommodate different applications like described

above.

In Fig. 3.4, by setting the PSO output to low voltage (changing jumper JP1 from

default 1-2 to 2-3) it was possible to make the laser turn off as a default and control

the laser on-time in the order of several microseconds. By altering the position of the

jumper, one can laser machine the back side of the sample by one to a few thousands

number of pulses even at high 1 MHz repetition rate.

Figure 3.3: Burst of the laser pulses generated with the high repetition rate MHz ultra-

short laser system using AOM (a) to produce 100% on/off laser beam modulation and

(b) to create an envelop of 80% duty cycle which is shown in dotted square wave [84].


Figure 3.4: Npaq Control Assembly showing various jumpers. The jumper JP1 was

changed from the default position of 1-2 to 2-3.

3.3 Purpose-built two-photon fluorescence imaging

setup

For in-situ real-time demonstration of laser interactions with materials, a two-photon

fluorescence imaging tool was in our laser system. This setup allowed us to take advan-

tage of the accurate diagnostic tool in the same setup as the laser fabrication normally

takes place. The two-photon fluorescence imaging setup is helpful in working with both

biological samples as well as photonic devices. Taking 2D or 3D images of the sample

before and/or after laser machining provides feedback and guidance for laser processing

of the sample.

3.3.1 Hardware of the fluorescence microscope system

In this part, various components of the purpose-built two-photon fluorescence microscope

will be described. In our experimental setup, a fiber-amplified laser (IMRA Jewel D-400-


VR) was used to provide a pulse duration of 300 fs at the wavelength of 1045 nm and the

repetition rates between 0.1 to 5 MHz and average power of 500 mW. Depending on the

sample type, either fundamental (1045 nm) or second harmonic (522 nm) laser would be

applied to the sample to trigger two-photon excitation.

Two-photon fluorescence imaging of the blood vessel and microspheres was done at

the fundamental wavelength, while green was applied for waveguide imaging. A 40X-0.65

numerical aperture (Nikon, CFI PL ACHRO 40X-A/0.65/0.57MM), infinity-corrected

objective lens was used for the major part of experiments reported here. For comparing

the resolution of the instrument fluorescence images were recorded of microspheres for

the following: 100× 0.9 NA dry lens (Nikon, BD plan 100X/0.9) and 100× 1.25-NA

oil-immersion lens (Nikon, plan 100X/1.25 oil) in addition to the 40X-0.65 lens. The

laser average power on the sample was adjusted to an appropriate range by applying

appropriate ND filters and by controlling the angle of the half-wave plate attenuator via

G-code. All the imaging has been recorded at 1 MHz repetition rate of the laser.

The actual beam displacement inside the sample is different from the displacement

of the lens (Fig. 3.5). The ratio of the two displacement as a function of the NA which

is the numerical aperture of the lens, n1 and n2 that are refractive indices of the two

mediums is given by Eq. (3.5).

d1d2

=n1n2

√1− (NA

n1)2√

1− (NAn2

)2(3.1)

where d1 and d2 are the lens displacement and actual focus displacement inside the

material, respectively.

The beam waist of the Gaussian beam (w) was calculated from [84,85]:

w =λ ·M2

π ·NA. (3.2)

Here, we have assumed the beam coming to the lens is collimated, and NA is the


Figure 3.5: Depth correction of refraction of the light according to the Snell’s law at the

interface of two materials with different refractive indices n1 and n2.

numerical aperture of the lens, λ is the laser wavelength, and M2 is the beam quality

factor.

The Rayleigh range (zR) of the Gaussian beam with the diffraction-limited assumption

can be defined as:

zR =π · w2 · n

λ. (3.3)

where n is the refractive index of the medium between lens and sample. The depth of

focus is twice of the Rayleigh range.

The ratio of d1/d2, the beam spot size, Rayleigh range, and the depth of focus are

summarized in Table 3.3.1 where n=1.589 was applied to Eq. (3.1) as the refractive index

of the microsphere.

As it is shown in Fig. 3.6, the fluorescence emission after passing through the dichroic

mirror (high reflection at 1045 nm or 522 nm and high transmission otherwise) was cou-

pled to a single mode fiber by a focusing lens (f = 10 mm) and then guided to an avalanche

photodiode (APD (Boston Electronics, id100-50MMF)). The fiber was mounted on a


Lens specifications d1/d2 Beam spot size [µm] Rayleigh range [µm] Depth of focus [µm]

Nikon, 40X 0.65 NA 0.5241 0.86 2.22 4.44

Nikon, 100X 0.9 NA 0.3329 0.62 1.15 2.31

Nikon, 100X 1.25 NA 0.7491 0.45 0.88 1.77

Table 3.1: Optical focusing parameters related to three lenses.

holder that could be precisely positioned in X, Y, and Z to then maximize coupling of

fluorescence signal. The dichroic mirror properties including transmission and reflection

wavelength range were chosen according to the incident laser wavelength. Two different

filters in the fluorescence signal path were used to block the incident laser beam either

in the green (Semrock, NF01-526U-25) or in the IR range (CVI laser SPF-950-1.00),

in order to improve the signal to noise ratio of the fluorescence signal over reflected or

scattered laser light that may enter the detector. The schematic of the purpose-built

two-photon fluorescence setup can be seen in Fig. 3.6.

After the optical signal was detected and converted to the electronic signal by the

APD, the electric signal entered the pulse counting electronics where Time Correlated

Single Photon Counting, TCSPC, (Becker and Hickl SPC-830) was used. In the TCSPC

method, it was possible to measure the arrival time of the single photon pulse precisely

and get the detected signal intensity in each specific time interval. Each time interval

was related to the pixel size by a scanning parameter such as the scan speed. Hence, one

can build the spatial intensity distribution in the form of a matrix where each element

of the matrix corresponds to one pixel. This information was transferred to a computer

including the SPC (Single Photon Counting) module. The intensity distribution of the

imaging area could be observed by means of the related image acquisition software.

Combination of several x-y image stacks then rendered a three dimensional image through

a TCSPC Laser Microscope software with user friendly environment developed within

the Labview interface. A block diagram representing the whole fluorescence imaging

system including both optical and electrical parts can be seen in Fig. 3.7. The optical


Figure 3.6: Two-photon fluorescence imaging setup. See text for detaile

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by Samira Karimelahi - University of Toronto T-Space...Samira Karimelahi Masters of Applied Sciences Graduate Department of Electrical and Computer Engineering University of Toronto